I'i,~ic Di~fIDA~~,hani'$
()·w
~
MA:RG
VICTO,lI
,,"',-.i,""
a
THE
'!'
~
i
'.
BASIC BIOMECHANICS
of the
MUSCULOSKELETAL
SYSTEM
Margareta Nordin, P.T., Dr. Sci.
Director. Occupational and Industrial Orthopaedic Center (OIOC)
Hospital fOf Joint Diseases Orthopaedic Institute
Mt. Sinai NYU Health
Program of Ergonomics and Biomechanics, New York University
Research Professor
Department of Orthopaedi" and Environmental Heallh Science
School of Medicine. New York University
New York. New York
Victor H. Frankel, M.D., Ph.D., KNO
President Emeritus
Hospital for Joint Diseases OrthopaedIC Institute
Professor of Orthopaedic Surgery
New York University School of Medicine
New York, New York
Dawn Leger, Ph.D., Developmental Editor
Kajsa Forssen, Illustrator
Angela Lis, P.T., M.A.. Editorial Assistant
$~~ LipPINCOIT WILLIAMS & WILKINS
•
A wohl'r! Klllwtr (ornp.lny
l'hil.ldclphi'l • B"llimore • New"ork . london
[lucllOS Ailes' Hong Kong· Sydney· T{lkyo
/
foreword
Mechanics and biology have always fascinated
humankind, The irnportance or understanding
the biomechanics of the musculoskeletal system
cannot be underestimated, Much a([entian has
been paid in recent years to genetic and biomoleClilar research, but the stlld~' of the mechanics
of structure and of the whole body s}'stcm is still
of immense importance. Musculoskeletal ailments arc among the most prevalent disorders in
the wodd and will continue to grow as the population ages.
Since the days when I first studied biomechanics in Sweden with Carl Hirsch, through my
years as an orthopaedic surgeon, teacher, and researcher, I have alway's emphasized combining
basic and applied research with clinical experiencc, This text represents my fifth effort to integrate biomechanical knowledge into clinical
training for patient carc. It is not a simple task
but by relating the basic concepts of biomechanics to everyday life, rehabilitation. orthopaedics.
traumatology, and patient care are greatly enhanced. Biomechanics is a multidisciplinary specialty, and so we have made a special effort to invite contributors from many disciplines so that
individuals from dilTerent fields may feel comFortable reading this book.
Together with an invaluable team, Margareta
Nordin and I have produced Ihis third edition of
Basic Biol1lechanics oFthe A'lusclt!o.\'keletal Systelll,
The new edition is shall1cncc! and improved
thanks to the input from the students and resi·
dents in orthopaedics Ihal during the past 10 years
have llsed the texl. This book is written for students and with a major input from students and
will hopefully be used to edunHe students and residents for many ~Iears to come. Although the basic
information contained in the book remains largely
unchanged, a considerable amount of extra information has been provided throughoul. ~Ve have
also made a special point to document with the
key references any significant changes in the field
of biomechanics and rehabilitation.
It has always been m)' interest lO bridge the
gap between engineering knowledge and clinical
carc and praclice, This book is written primarily
for clinicians such as orthopaedists, physical
and occupational therapists, clinical ergonomisls, chiropractors, and olher health professionals who arc acquil-ing a working knowledge
of biomechanical principles for use in the evaluation and treatment of musculoskeletal dysfunction. We only hope that if you find this book int~resting. you will seek more in-depth study in
the field of biomechanics. Enjo\' it. discuss it,
and become a beller clinician and/or researchcl:
Vve are extremely proud that Basic BiomeclUluics oj" the !\tlllscliioskeic/lli Sysle111 has
been designated "A Classic" by the publishers,
Lippincott Williams & Wilkins. We Ihank the
readers, students, professors, and all who acquire thc tcxt and lise it.
VielO,. H. Frallkel, M.D., Ph.D., KNO
vii
l'~ e"!,-g,! !,•.-~--~"".'!~~~.""~"""c""',:::~::""'-'''''''';'-:'''"''''.''''''·'''''--'''''--''''''''''':~'''i~!1'···"'-"-"'·E"'.r,;>d"._"".""
.....,. ~"-h_,,:_/>;:~B:>~'¥,,..'13iA:s"' *w.c-:;·_!l , _~_iIl~!il i.iIl=i1!~i1!i!i!lWl;41.4~l'l!~$~rm!1il·, ~·" "'1l i&~.l\l.l!l~t!%_"t81; .
.,
Preface
ever possible. We retained the selected examples
to illustrate lhe concepts needed for basic knowledge of the musculoskeletal biomechanics; we
also have kept the important engineering concepts throughout the volume. We have added
four chapters on applied biomechanics topics.
Patient case studies ancl calculation bo:'\cs have
been added to each chapter. \Ne incorporated
flowcharts throughout the book as teaching tools.
The text will serve as guide to a deeper understanding of musculoskeletal biomcchanics gained
through funher reading and independent research.
The information presented should also guide the
rt.'ader in assessing the literature on biomechanics.
"Vc have attemptcd to provide therapcutic examples but it was not our purpose to cover this area;
instead, \ve have described the undel'lying basis for
rational therapcutic or exercise programs.
An introductory chapter describes the inlporlance of the study of biomechanics, and an appendix on the international system of measurements serves as an introduction to the physical
measurements used throughout the book. The
reader needs no more (han basic knowledge of
mathematics to fully comprehend the material
in the book, but it is important to review the appendix on the Sl System and its application to
biomechanics.
The body of the third edition is then divided
into three sections. The first section is the Biomechanics of Tissues and Stnlcturcs of the Musculoskeletal System and covers the basic biomechanics of bone, ligaments, cartilage. tendons,
muscles, and nenres. The second section covers
the Biomechanics-of Joints, including every joint
system in the human body. Chapters range from
the foot and ankle through the cervical spine,
and co\'er eve I:" joint in between. The third sec-
Biomechanics uses physics and engineering concepts to describe the motion undergone by the
various body segments and forces acting on
these body parts during normal activities. The
inter-relationship of force and motion is important and must be understood if rational treatment programs are to be applied to musculoskeletal disorders. Deleterious effects may be
produced if the forces acting on the areas with
disorders rise 10 high levels during exercise or
other activities of daily living.
The purpose of this text is to acquaint the readers with the force-motion relationship \vithin the
musculoskeletal system and the various tech~
niqucs llsed to understand these relationships.
The third edition of Basic Biol1/eclwllics of rite
iHllScliloskeleral System is intended for use as a
textbook either in conjunction with an introductory biomechanics course or for independent
study. The third edition has been changed in
many ways, but it is still a book that is designed
for use by students who are interested in and
want to learn about biomechanics. It is primarily
written for students who do not have an engineering background but who want to understand
the most basic concepts in biomechanics and
physics and how these apply to the human body
Input from students has greatly improved this
third edition. We have used the book for 10 years
in the Program of Ergonomics and Biomechanics
at New York University', and it is the students and
residents who have suggested the changes and
who have continuously shown an interest in developing and irnproving this book. This edition
has been further strengthened by the contribution or the students over the past year. \Vc formed
focus groups to understand better what thc students wanted and applied their suggestions wher-
e .-..
---~-_.--
. _--.-._.. _------- .-.- . ----.. -.-....
-'------------~---
..
,~".--
. . -.. ..- ...-.--..-------..
~.~
~
-.,,"------~-----
-",,--.-.---.------
~
tion cover~ some topics in Applied Biomechanics, including chapters on fracture fixation;
arthroplasty; silting, standing and lying; and
gait. These arc basic chapters that sl:l\r e to intra·
c1uce topics in applied biomechanics: they arc
not in-depth explorations of the subject.
Finally. we hope that the revision and expan-
oFthe kluscu/oskelela/ Syslel1l will bring about an
increased awareness of the imparlance of biomechanics. II has never been our intention to
complL'tely cover the subject, but instead provide
a basic introduction to the field that will lead to
further study or this important lopic.
sion of this third edition of" Basic 13io11leclulJIics
Margarela NOf(!;11 alld Viclor H. Frankel
""at ~.,,;,~~~~ :,,~,~t,
Acknowledgments
_--..
---------~
I
i
!
!Ii
I
!
!
I
I
;
This book was made possible through the outstanding contributions of many individuals. The
chapter authors' knowledge and understanding of
the basic concepts of biomechanics and their
wealth of experience have brought both breadth
and depth to this work. Over the past 10 years.
questions raised by students and residents have
made this book a better teaching tool. The Third
Edition could not have been done without the
students who have shared their cornmen(s and
really sCnItinizcd thc Second Edition. There arc
too many names LO list here, but we thank each
student who asked a question or made a suggestion during the course of his or her studies.
Special thanks to the students who panicipated
in several focus groups. whose input was invaluable in finalizing the contents and design of
the text.
Vve are honored and grateful for the contributions of everyone who has worked to prepare
this new edition. 'vVe can honestly say that this
third edition is written ror the sludent and by
students and residents who leave the classroom
with the knowledge to enhance our life and
existence.
A book of this size with its large number of
figures, legends, and references cannot be produced without a strong editorial team. As project
editor, Dawn Leger's continuous effort and
perseverance and thoughtfulness shines through
the entire book. She has contributed not just to
the editing but also to logistics, and as a stylist,
as an innovator, and a friend. Our editorial
assistant, Angela Lis, is a physical thcrapisl and
recent recipient of the MA degree in Ergonomics
and Biomechanics from NYU. As a recent graduate, Angela was also a recent USCI' of the book,
and she devoted several months to help finalize
this edition. She created the flowcharts and scrutinized all the figures, patient cases, and caku-
-1.'- •
JUlian boxes. Angela look this book to her hear
and \ve arc all the bettcr for her passion an
attention to detail.
The illustrator: Kajsa Forssen. has now worke
on all three editions of this text. Her never-failin
grasp of hiomechanical illustrations, her simp
city and exactness of figures, is always appr
ciated. In drawing all the figures and graphs, sh
considers how they would translate into a slide
into a computer-generated presentation. Kaj
Forssen is one of the top iIIustralOrs that we hav
ever worked wilh, and she has been an importa
member of the publication (cam.
This book was also made when publicatio
companies I11ergcd and merged again, and in th
end we are deeply grateful to Ulila Lushnyck
who has with her team at Lippincott \·Villiams
Wilkins been responsible ror the production. Sh
has worked with tremendous energy and posiliv
thinking, put the book together in record spce
and we fonvard our sincerest gratitude to he
\Ne are also thankful for a development gra
provided by Lippincott Williams & Wilkins
finance this effort.
Our colleagues al the Occupational an
Industrial Orthopaedic Cenler and the Depar
ment or Orthopaedics of the Hospital ror Joi
Diseases Orthopaedic Institule functioned
critical reviewers and contributors to th
chapters. Special thanks is extended to Dav
Goldsheydcl" for assislance in reviewing lh
biomechanical calculation boxes. to Marc
Campello as a contributor and reviewer, and
Shira Schccter-vVeiner for contributing to th
spine chapteI: Much thanks to Dr. Mark Pitma
1'01' supplying vital x-rays ror the new edition. \,V
are parlicularly grateful to DI: Markus Pielr
for contributing with the latcst on intr
abdominal pressure. to Dr. Ali Sheikhzadeh r
reviewing chapters and contributing ne
•
references, to Dr. Tobias Lorenz for his work on
the first section, and to all other staff at the
Occupational and Industrial Orthopaedic Center
who have been managing the center while we are
absorbed wilh the book.
\'\Fe arc most grateful to Drs. Bejjani, Lindh,
Pitman, Peterson, and Stuchin for their COI1l1·j·
bUlions to the second cdition which sen'cd as a
framework for the updated third edition.
The third edition of Basic Biomechallics orEiIe
iHuscllloskeletal System was supported throughout its production by the Research and Development Foundation of the Hospital for Joint
Diseases Orthopaedic Institute and the hospital
administration, to whom we forward our sincere
gratitude.
To all who helped, we say' again, thank yOLi
and TACK SA MYCKET.
klargareta Nordin and Victor fl. Frankel
Contributors
Gunnar B. J. Andersson, M.D., Ph.D.
Professor and Chairman
Department of Orthopaedic Surgery
Rush-Presbyterian-SI. Luke's Medical Center
Chicago, IL
Thomas P. Andriacchi, Ph.D.
Biomechanical Engineering Division
Stanford University
Stanford, CA
Victor H. Frankel, M.D., Ph.D., KNO
President Emeritus
Sherry I. Backus, M.D., P.T.
Senior Research Physical Therapist and Research Associate
Motion Analysis laboratory
Hospital for Special Surgery
New York, NY
Ann E. Barr, Ph.D., P.T.
Physical Therapy Department
College of Allied Health Professionals
Temple University
Philadelpllla, PA
Fadi Joseph Bejjani. M.D .. Ph.D.
Director of Occupational Musculoskeletal Diseases
Department
University Rehabilitation Association
Newark, NJ
Maureen Gallagher Birdzell, Ph.D.
Departmenl of Orthopaedic Surgery
Hospital for Joint DiseasesiMI. Sinai NYU Health
New York, NY
Marco Campello, P.T., M.A.
Associate Clinical Director
Occupational and Industrial Orthopaedic Center
Hospital for Joint DiseasesiMI. Sinai NYU Health
New York. NY
Dennis R. Carter. Ph.D.
Professor
Biomechanical Engineering Program
Stanford University
Stanford, CA
-_.
__._---._._-_._--_._
Hospital for Joint Diseases Orthopaedic Institute
Professor of Orthopaedic Surgery
New York University School of Medicine
NevI York, NY
Ross Todd Hockenbury, M.D.
River City Ortl1opaedic Surgeons
LouiSVille, KY
Assistant Professor
~.-
Craig J, Della Valle, M.D.
NYU-HJD Department of Orthopaedic Surgery
Hospital for Joint Diseases
School of Medicine
New York University
New York, NY
_._._.-
Clark T. Hung, Ph.D.
Assistant Professor
Department of Mechanical Engineering and Center for
Biomedical Engineering
Columbia University
New York, NY
Debra E. Hurwitz. Ph.D.
Assistant Professor
Department of Orthopaedics
Rush·Presbyterian-St. Luke's lvIedical Center
Chicago, IL
Laith M. Jazrawi. M.D.
NYU·HJD Department of Orthopaedic Surgery
Hospital for Joint Diseases
School of Medicine
New York University
New York, NY
Frederick J, Kummer, Ph.D.
Associate Director, Musculoskeletal Research Center
Hospital for Joint DiseasesiMt. Sinai NYU Health
Research Professor, NYU-HJD Department of Orthopaedic
Surgery
Scl100l of Medicine
New York University
New York, NY
_._-_._._._-_. _.-.._.-._-_ _._._-_._ _ _-_._-- _
__ _._.
xiii
Dawn Leger, Ph.D.
Adjunct Assistant Professor
NYU-HJD Department of Orthopaedics
School of Medicine
New York University
New York, NY
Jane Bear-Lehman, Ph.D., OTR, FAOTA
Assistant Professor of Clinical Occupational Therapy
Department of Occupational Therapy
Columbia University College of Physicians and Surgeons
New York. NY
Margareta Lindh, M.D., Ph.D.
Associate Professor
Department of Physical MeeJicine and Rehabilitation
Sahlgren Hospital
Gothenburg University
Gothenburg, Sweden
Angela Lis, M.A., P.T.
Research Physical Therapist
Occupational and Industrial Orthopaedic Center
Hospital for Joint DiseasesiMt. Sinai NYU Health
New York, NY
Associate Professor
Physical Therapy Program
(orporacion Universitaria Iberoamericana
Bogota, COLOMBIA
Tobias Lorenz, M.D.
Fellow
Occupational and Industrial Orthopaedic Center
Hospital for Joint Diseases/Me Sinai NYU Health
New York, NY
Goran Lundborg, M.D.
Professor
Department of Hand Surgery
Lunds University
Malmo Allmanna Sjukhus
Malmo, Sweden
Ronald Moskovich, M.D.
Associate Chief
Spine Surgery
NYU-HJD Department oj Orthopaedic Surgery
Hospital for Joint Diseases
School of Medicine
New York University
New York, NY
Van C. Mow, Ph.D.
Director
Orthopaedic Research Laboratory
Department of Orthopaedic Surgery
Columbia University
New York, NY
Robert R. Myers, Ph.D.
Associate Professor
Department of Anesthesiology
University of California San Diego
La Jolla, CA
Margareta Nordin, P.T., Dr. Sci.
Director, Occupational and Industrial Orthopaedic Center
(OIOC)
Hospital for Joint Diseases Orthopaedic Institute
!vlt. Sinal NYU Health
Program of Ergonomics and Biomechanics
New York University
Research Professor
Department of Orthopaedics and Environmental Health
Science
School of Medicine, New York University
New York, NY
Kjell Olmarker, M.D., Ph.D.
Associate Professor
Department of Orthopaedics
Sahlgren Hospital
Gothenburg University
Gothenburg, Sweden
Nihat bzkaya (deceased)
Associate Professor
Occupational and Industrial Orthopaedic Center
Hospital for Joint Diseases
Research Associate Professor
Department of Environmental Medicine
New York University
NelN York, NY
Lars Peterson, M.D., Ph.D.
Gruvgat 6
Vastra Frolunda
Sweden
Mark I. Pitman, M.D.
Clinical Associate Professor
NYU-HJD Department of Orthopaedic Surgery
School of lvIedicine
New York University
New York, NY
Andrew S. Rokito, M.D.
Associate Chief. Spons Medicine Service
ASSistant Professor
NYU-HJD Department 'Of Ortllopaedic Surgery
School of Medicine
New York University
New York, NY
Bjorn Rydevik, M.D., Ph.D.
Professor and Chairman
Department of Orthopaedics
Sahlgren Hospital
Gothenburg University
Gothenburg. Sweden
G. James Sammarco, M.D.
Program Director
fellowship in Adult Reconstructive Surgery
foot and Ankle Orthopaedic Surgery Program
The Center for Orthopaedic Care, Inc.
Volunteer Professor of Orthopaedic Surgery
Department of Orthopaedics
University of Cincinnati Medical Center
Cincinnati, OH
Chris J. Snijders, Ph.D.
Professor
Biomedical Physics and Technology
faculty of Medicine
Erasmus University
Rotterdam, The Netherlands
Steven Stuchin, M,D.
Director Clinical Orthopaedic Services
Director Arthritis SefYice
Associale Professor
NYU-HJD Department of Orthopaedics
School of Medicine
New York University
New York, NY
Shira Schecter Weiner, M.A., P.T.
Research Physical Therapist
Occupational and Industrial Orthopaedic Center
Hospital for Joint Diseases/Mt. Sinai NYU Health
New York. NY
Joseph D. Zuckerman, M.D.
Professor and Chairman
NYU-HJD Department of Orthopaedic Surgery
Hospital for Joint Diseases
School of Medicine
New York University
New York, NY
Contents
Introduction to Biomechanics: Basic
Terminology and Concepts 2
Biomechanics of the Foot
and Ankle 222
Niha! bzkaya, Dawn Leger
G. James Sammarco, Ross Todd Hockenbury
Appendix 1: The System International
d'Unites (SI) 18
(11)
Biomechanics of Bone
2
Margareta Nordin, Shira Schecter Weiner,
adapted from Margareta Lindh
Dennis R. Carter
trw
~
Biomechanics of Tissues
and Structures
of the Musculoskeletal System
Biomechanics of the Lumbar Spine
Biomechanics of the Cervical Spine
2
Ronald Moskovich
(f}
Biomechanics of the Shoulder
318
Craig J. Della Valle. Andrew S. Rokito. Mauree
Gallagher Birdzell, Joseph D. Zuckerman
Biomechanics of the Elbow
26
340
Laith M. Jazrawi, Andrew S. Rokito, Maureen
Gallagher Birdzell, Joseph D. Zuckerman
Victor H. Frankel, Margareta Nordin
Biomechanics of Articular
Cartilage 60
Biomechanics of the Wrist
and Hand 358
Van C. Mow, Clark T. Hung
Ann E. Barr, Jane Bear-lehman adapted from
Steven Stuchin, Fadi J. Bejjani
Biomechanics of Tendons
and Ligaments 102
Margareta Nordin, Tobias Lorenz, Marco
Campello
e
Applied Biomechanics
Biomechanics of Peripheral Nerves
and Spinal Nerve Roots 126
Bjorn Rydevik, Goran lundborg, Kjell Olmarker,
Robert R. Myers
Biomechanics of Skeletal Muscle
~ Introduction to the Biomechanics
of Fracture Fixation 390
Frederick J. Kummer
148
Biomechanics of Arthroplasty
Tobias lorenz, Marco Campell0
adapted from Mark 1. Pitman, Lars Peterson
([$
~.. ,
Biomechanics of Joints
e
o
Biomechanics of the Knee
400
Debra E. Hurwitz, Thomas P. Andriacchi, Gunn
B.J. Andersson
Engineering Approaches to Standing,
Sitting, and Lying 420
Chris J. Snijders
176
ED Biomechanics of Gait
Margareta Nordin, Vietor H. Frankel
Ann E. Barr, Sherry l. Backus
Biomechanics of the Hip
Index
202
438
459
Margareta Nordin, Victor H. Frankel
x
BASIC BIOMECHANIC
of the
MUSCULOSKELETA
SYSTE
-~
'"
I
Introduction to
Biomechanics:
Basic Terminology
and Concepts
Nihat OZkaya, Dawn Leger
Introduction
Basic Concepts
Scalars, VeCtors, and Tensors
Force Vector
Torque and Moment Vectors
Newton '$ l.aws
Free-Body Diagrams
Conditions for Equilibrium
Statics
Modes of Deformation
Normal and Shear Stresses
Normal and Shear Strains
Shl?ar·5train Diagrams
Elastic and Plastic Deformations
Viscoelasticity
Material Properties Based 011 Stress-Strain Diagrams
Principal Stresses
Fatigue and Endurance
Basic Biomechanics of the Musculoskeletal System
Part I: Biomechanics of Tissues and Structures
Part 11: Biomechanics of Joints
Part III: Applied Biomechanics
Summary
Suggested Reading
Introduction
Biomechanics is considered a branch of bioengineering and biomedical engineering. Bioengineering is an interdisciplinary field in which the principles and methods from engineering. basic sciences.
and technology arc applied to design. test, and manufacture equipment for use in medicine and to understand, define, and solve problems in physiology
and biology!. Bioengineering is one of several specialty areas that corne under the general field of biomedical engineering.
Biomechanics considers the applications of classical mechanics 10 the analysis of biological and
physiological svstems. Different aspects of biomechanics utilize different parts or applied mechanics.
For example, the principles of statics havc been applied to analyze the magnitude and nature of forces
involved in various joints and muscles of the nUlsculoskeletal system. The principles of dynamics
have been utilized for motion description, gait
analysis, and segmental motion analysis and have
many applications in sports mechanics. Thc mc~
chanics of solids provides the necessary tools for
developing the field constitutive equations For biological systems that are used to evaluate their functional behavior under dilTerent load conditions. The
principles of fluid mechanics have been used to investigate blood flow in the circulatory system, air
flow in the lung, and joint lubl'ication.
Research in biomechanics is aimed at improving
our knowledge of a vcry complex structure-the human body. Research activities in biomechanics can
be divided into three areas: experimcntal studies,
model analyscs, and applied research. Experimental
studies in biomechanics arc done to determine the
mechanical properties of biological materials, in~
eluding the bone, cartilage, muscle, tendon, ligament. skin, and blood as a whole or as parts
constituting them. Theoretical studies involving
mathematical model analy1ses have also been an im~
ponant component of research in biomechanics. In
general, a model that is based on experimental findings can be used to predict the efrect of environmental and operational factors without resorting to
laboratory experiments.
Applied research in biomechanics is the application of scientific knowledge to bcnefit human bcings. vVe know that musculoskeletal injury and illness is one of the primary occupational hazards in
industrialized countries. By learning how the musculoskeletal system adjusts to common work conclitions and by developing guidelines to assure that
manual work conforms more closely to rhe physica
limitations of the human body and to natural bod
rnO\'cmCnlS, these injuries rnay be combatlcd.
Basic Concepts
Biomechanics of the musculoskeletal system r
quires a good understanding of basic mechanic
The basic terminology and concepts from mechan
ics and physics arc utilized to clcscribe intcrn
forces of the human body. The objective of studyin
thcs~ forces is to understand the loading conditio
of soft tissues and their mechanical responses. Th
purpose of this section is to rC\'jew the basic con
cepts of applied mechanics that are used in biome
chanical literature and throughout this book.
SCALARS, VECTORS, AND TENSORS
Most of the concepts in mechanics arc either scal
or vector. A scalar quanlity has a magnitude onl
Concepts such as mass, energy', power, mechanic
work, and temperalure are scalar quantities. For e
ample, it is suffkicnt to say that an object has 8
kilograms (kg) of mass. A vector quanlity, con
versely, has both a magnitude and a direction ass
ciated \vith it. Force, moment, velOcity, and accele
ation arc exall'lples of vector quanlities. To describ
a force fully. one must state how much force is a
plied and in which direction it is applied. The ma
nitude of a vector is also a scalar quantity. The ma
nitude of any quanlity (scalar or vector) is always
positi\'c number corresponding to the numeric
measure of that quantity_
Graphically, a vector is represented by an arrow
The orientation of the alTow indicates the line of a
tion and the arrowhead denotes the direction an
sensc of the vectm: 'If 1110re than one vector must b
shown in a single drawing, the length of each arro
must be proportional to the Inagnitude of the vect
it represents. Both scalars and vectors arc speci
forms of a more general category of all quantities
mechanics called tensors. Scalars arc also known
"zero-01"(Ic1· tensors," whereas vectors aJ'e "firs
order tensors." Concepts such as stress and strai
conversely, are "second-order tensors."
FORCE VECTOR
Force can be defined as· mechanical disturbance
load. Whcn an object is pushed or pulled, a force
applied on it. A force is also applied when a ball
thrown or kicked. A force <:\Cling on an object may
deform the objecl, change its slale of m~lion, 0"1'
both. Forces may be c1assif-lcd in variolls ways according to their effects on the objects to \vhicf'l they
arc applied or according to their orientation as con~­
pared with one another. For example, a force may
be internal or external, normal (perpendicular) 0'1'
tangential; tensile. compressive. 01" shear; gravitational (weight); or frictional. Any two or more forces
acting on (\ single body 111ay be coplanar (acting on
a lwo~dimcnsional plane surface); collinear (have a
common line of action); concurrent (lines of action
intersecting at a single point); or parallel. Note that
weight is a special form of Force. The weight of an
object on Earth is the gravitational force exerted bv
Earlh on the mass of that object. Thc magnitude ~'r
the wcighl of an object on Ea~·t11 is equal
the mass
of the object times the magnitude of the gravitational acceleration, \vhich is approximately 9.8 mt>
ters pCI' second squared (111/s 1). For exampl~, a 10-k<J
object weighs approximately 98 newtons (N) o~
Earth. The direction of weight is always vertically
do\vl1\vard.
t;
TORQUE AND MOMENT VECTORS
The effect of a roree on the object it is applied upon
depends on how the rorce is applied and how the
object is suppo!"ted. For example, when pulled. an
open door will swing about the edge along which it
is hinged lO the wall. \-Vh'll eallses the door 10 swin cJ
is the torque generated by the applied force abou~
an axis that passes through the hinges of the door. If
one stands on the free-end of a diving board, the
hoard will bend, What bends the board is the momel1l of the body weight about the fixed end of the
board. In general. torque is associated with the ro~
tnlional and twisting action of applied forces, while
moment is related to the bending action. However,
the mathematical defmition of moment and torque
is the same.
Torque and moment arc vector quantities. The
magnitude of the tonlue Of rnoment of a force
about a point is equal to the mannitude of the
force times the length of the shortc:t distance between the point and the line of action of the force
which is known as the lever or moment arm. Con~
W)-_-I
Definition of torque. Reprinred with permission from DZkaya, N. (998). Biomechanics.
In w.N. Rom, Environmental and Occupationa( Medicine (3rd ed., pp, 1437~1454), New
York: Lippincott·Raven,
sider a person on an exercise apparatLls who is
holding a handle that is attached to a cable (Fig.
I-I), The cable is wrapped around a pulley and attached to a weight pan. The weight in the \veight
pan stretches the cable such that the magnitude F
of the tensile force in the cable is equal 10 thc
weight of the weight pan. This force is transmitted
to the person's hand through the handle, At this instant, if the cable allached to the handle makes an
angle 0 with the horizontal. then the force E exerted by the cable on the person's hand also makes
an angle 0 with the horizontal. Let 0 be a point on
the axis of rolation of the elbow joint. To dctermine the magnitude of the moment due to force f
about 0, extend the line of action of force f and
drop a line from 0 that cuts the line of action of F
at right angles. If the point of intersection or the
twO lines is Q, then the dbtance d between 0 and
Q is the lever arm, and thc magnitude of the rno~
ment M of force E about the clbow joint is M = dE
The direction of the moment ,·cctor is perpendicular to thc plane defined by the line of action of E
and line 00, or for thb two-dimensional case, it is
counterclockwise.
NEWTON'S LAWS
Relatively few basic laws govern the relationship
betwcen applied forces and corresponding motions, Among these, the laws of mechanics introduced by Sir Isaac Newton (1642-1727) are the
most important. NeWLOn's first law states that an
object at rest will remain at rest 01' an object in motion will move in a straight line with constant velocity if the net force acting on the object is zero.
Newton's second law states that an object with a
nonzero net force acting on it will accelerate in the
direction of the net force and that the magnitude of
the acceleration will be proportional to the magnitude of the net force. Newton's sccond law can be
formulated as E = m ;), Here, E is the applied force.
m is the mass of the object, and i! is the linear
(translational) accelcration of the object on which
the force is applied. If more than one force is acting
on the object. then E represcnts the net or the resultant force (the vector sum of all forces), Another
way of stating Newton's second law of motion is
M = I Q., where M is the net or resultant moment of
all forces acting on the objcct, I is thc mass moment
of inertia of the object, and ~ is the angular (rota- ,
tional) acceleration of the object. The mass m and
mass moment of incrtia I in these equations of motion arc measures of resistance to changes in 1110-
lion. Th~ larger the inertia of an object, the m
difficult it is to sel in motion or to SlOp if it is
rend)' in motion.
Newlon's third law states that to every act
there is a reaction and that the forces of action a
reaction between interacting objects are equal
magnitude, oppositc in direction, and have
same line of action. This law has important appli
tions in constructing free· body diagrams.
FREE· BODY DIAGRAMS
Free-body diagrams are constructed to help iden
the forces and moments acting on individual pa
of a system and to ensure the correct use of
equations of mechanics La analyze the system.
this purpose. the parts constituting a system are i
lated from their surroundings and the effects of s
roundings arc replaced by proper forces and m
ments.
The human musculoskeletal system consists
many parts that are connected to one anot
through a cornplcx tendon. Iigamcnt, muscle, a
joint SU·uctUfC. In somc analyses, the objective m
be to investigate the forces involved at and arou
val'ious joints of the human body for different p
tural and load conditions. Such analyses can be c
ried out by separating the body into two parts al
joint of interest and drawing the free-body diagr
of one of the parts. For example, consider the a
illustrated ill Figure I ~2. Assume thal the forces
volved at the elbow joint arc to be analyzed. As
lustrated in Figure 1-2, lhe entire body is separa
into two at the elbow joint and the free-body d
gram of the forearm is drawn (Fig. 1-2B). Here,
E is the force applied to the hand by the handle
the cable attached to the weight in the weight pa
\V is the total wcight of thc lower arm acting
the center of gravity of the lower arm,
£,\\1 is the force excrted by' the biceps on the
dius,
£.,,; is the force exerted bv the brachioradi
muscles on the radius.
£.\l~ is the force exerted by the brachialis musc
on the ulna, and
f1 is the resultarll reaction force at the hume
ulnar and humeroradial joints of the elbow. N
that the muscle and joint reaction forces repres
the mechanical effects of the upper ann on
lower arm. Also note that as illustrated in Fig
1-2;\ (which is not a complete free-body diagra
equal magnitudc but opposite muscle and joint
action forces act on the upper arm as wcll.
"
-,
'
.~
...
dum implies that the body of concern is either at
rest or moving with constant velocity. For a body to
be in a slate of equilibrium, it has to be both in
translational and rotational cquilibl-iul11. A body is
in translational cquilibriun1 if the net force (vector
sum of all forces) acting on it is zero. If the Ilt:t force
is zero, then the linear acceleration (time rate of
change of linear velocity) of the body is zero, or the
linear velocity of the bod,Y is either constant or zero.
A body is in rotational equilibrium if the net moment (vector sum of the moments of all forces) act·
ing on it is zero. If the net moment is zero, then the
angular acceleration (time rate of change of angular
velocity) of the body is zero, or the angulal· yelocily
of the body is either constant or zero. Therefore, for
a body in a state of equilibrium, the equations of
motion (Newton's second law) take the following
special forms:
A
~E =
E
B
w
Forces involved at and around the elbow joint
and the free-body diagram of the lower arm.
Reprinted with permission from dzkaya. N.
(7998). Biomechanics. In W.N. Rom, Environmental and Occupational Medicine (31d ed., pp.
1437-1454). New York: Lippincott-Raven.
CONDITIONS FOR EQUILIBRIUM
Statics is an area within applied mechanics that is
concerned with the anal~:sis of forces on rigid bodies in equilibrium. A rigid body is one that is assumed to undergo no deformations. [n reality, evcr)'
object or matcrial may undergo deformat ion to an
extent when acted on by forces. (n some cases, the
amount of deformation may be so smalllhal il may
not affect the desired analysis and the object is assumed to be rigid. In mechanics, the term cquilib-
0 and
~rvl
=0
rt is important to remember that force and moment arc vector quantities. For example, with respect to a rectangular (Cartesian) coordinate system, force and moment vectors may have
components in the .'\. y, and z directions. Therefore,
if the net force acting on an object is zero, then the
sum of forces acting in each direction must be equal
lo zero (IF, = 0, IF, = 0, IF, = 0). Similarly, if the
net moment on an object is zero. then the sum of
moments in each direction must also be equal to
zero (lM, = 0, lM,. = 0, lM, = 0). Thel'efore, for
three-dimension force systems there arc six cOl1ciilions of equilibrium. For two-dimensional force sys[ems in lhe xy-plane, onl~: three of these conditions
(IF, = 0, ~F, = 0, and ~M, = 0) need to be checked.
STATICS
The principles of slatics (equations of equilibrium)
can be applied to investigate the muscle and joint
forces involved at and around the joints for various
postural positions of the human body and its segments. The immediate purpose of static analysis is
to provide answers to questions such as: What tension must the neck extensor muscles exert on the
head to support the head in a specined position?
\OVhen a person bends, what would be the force ex~
ertcd by the erector spinae on the fifth lumbar vertebra? Ho\\'" does the con1pression at the elbow,
knee, and ankle joints vmy with externally applied
forces and with different segmental arrangements?
How docs the force on the femoral head vary with
loads carried in the hand? \,Vhat arc the forces in~
volved in various muscle groups and joints during
different exercise conditions?
In general. the unknowns in static problems involving the musculoskeletal s:'stcm arc thc magnitudes of joint reaction forces and muscle tensions.
The mechanical analysis of a skelclal joint requires
that we know the vector characteristics of tensions
in the muscles, the proper locations of muscle attachments, the weights of body segmcnts, and the
locations of the centers of gravity of the body segmems. Mechanical models are obviously simple
representations of c0l11plex systems. Many models
are limited by the assumptions that must be made
to reduce the system under considcration to a st.atically determinate one. Anv model can be improved
by ~onsidering the comril)utions of other muscles,
but (hat will increase the number of unknowns and
make the model a statically indeterminate one. To
analyze the improved model. the researcher would
need additional information related to the muscle
forces. This inforrrlalion can be gathered through
electromyography measurements of muscle signals
or by applying certain optirnization techniques. A
similar analysis can be made to investigate forces
involved at and around other major joints of the
musculoskeletal system.
MODES OF DEFORMATION
When acted on by externally applied forces. objects
may translate in the direction of the net force and
rotate in the direction of the net torque acting on
them. If an object is subjected to externally applied
forces but is in stalic equilibrium. then it is most
likely that there is some local shape change within
the objec!. Local shape change under the effect of
applied forces is known as deformation. The extent
of deformation an object may undergo depends on
many' factors, including the material properties.
size, and shape of the object; environmental factors
such as heat and humidity; and the nlagnitudc, direction, and duration of applied forces.
One way of distinguishing forces is by observing
their tendencv to deform the object they are applied
upon. For example. the object is said to be in tension if the body tends to elongate and in compression if it tends to shrink in the direction of the applied forces. Shear loading differs from tension and
compression in that it is caused b:! forces acting in
directions tangent to the area resisting the forces ~
causing shear, whereas both tension and compression are caused by collinear forces applied perpendicular to the areas on which they act. It is common
':,
to call tensile and compressive forces normal o
ial forces: shearing forces are tangential forces
jects also deform when they are subjecled to f
that cause bending and torsion, which are relat
the moment and torque actions of applied forc
A matel"ial nwv respond differently to diff
loading configurations. For a given material.
may be different physical properties that mu
considered while analyzing the response of tha
terial to tensile loading as compared with com
sive or sheai' loading, The mechanical propert
mnterials are established through stress analys
subjecting them to various experiments such as
axial tension and compression, torsion, and
ing tests.
NORMAL AND SHEAR STRESSES
Consider the whole bone in Figure \-3;,\ that is
jected to a pair of tensile forces of magnitude F
bone is in static equilibriulll. To analyze the f
induced within the bone, the method of section
be applied by hypothetically cutting the bone
two pieces through a plane perpendicular to the
axis or the bone. Because the bone as a whole
equilibrium, the two pieces must individually
equilibrium as well. This requires that at the cu
tion of each piece there is an internal force t
equal in magnitude but opposite in direction t
externally applied force (Fig. 1-38). The int
force is distributed over the entire cross-sec
area of the cut section. and E represents the resu
of the distributed force (Fig. 1-3C). The intens
this distributed force (force per unit area) is k
as stress. For the case shown in Figure 1-3. be
the force resultant at the cut section is perpendi
to the plane of the cut. the cOITesponciing str
called a normal or axial stress. It is customar:y t
the symbol (T (sigma) to refer to normal stresse
suming that the intensity of the distributed fo
the Cllt section is uniform over the cross-sec
area A of the bone, then u::::: FlA. Normal stresse
are caused by forces that tend to stretch (elon
matcl"ials aJ"C marc specincally known as t
stresses; those that tend to shrink them are kno
compressive stresses. According to the Standar
ternational (SO unit system (see Appendix), str
are measured in newton per square meter (N
which is also known as pascal (Pa).
There is another form of stress, shear s
which is a measure of the intensity of internal f
acting tangent (parallel) to a plane of cut
~i
H'
r
'.i~
,..-__..J-_=-
-
F .....~--I
..
~-
F
A
A
~-G?-
F ......
=' ~,
B
c
Definition of normal stress. Reprinted wirh permission from OZkaya. N. (1998j. Biomechanics. In W.N. Rom, Environmental and Occupatiol1<11 r..,ledicine (3rd ed., pp.
i437-145r+). New York: Lippincorr·RiNen.
,
d;
!:
example. consider the whole bone in Figure 1-4A.
The bone is subject to a number of parallel forces
that act in planes perpendicular to the long a,is of
the bone. Assume that the bone is cut into two parts
through a plane perpendicular to the long axis of
the bone (Fig. 1-48). If the bone as a whole is in
equilibrium, its individual parts must be in equilibrillm as well. This requires that there must be an internal force at the cut section that ,lets in a direction
tangent to the cut surFace. If the magnitudes of the
external forces arc known, then the magnitude F of
the internal force can be calculated by considering
the translational and rotational equilibrium of onc
of the parts constituting the bone. The intensity of
the internal force tangent to the Clit section is
known as the shear stress. It is customary to usc the
symbol T (tau) to refer to shear stresses (Fig. 1-4C).
Assuming that the intensity of the force tangent to
the cut section is uniform over the cross-sectional
area A of the bone, then T = FlA.
NORMAL AND SHEAR STRAINS
Strain is a measure of the degree of deformation. As
in the case of stress, two types of strains can be distinguished. A norm~l'l strain is deflnecl as the ratio of
the change (increase or decrease) in length to the
original (undeformed) length, and is commonly denoted with the symbol € (epsilon). Consider the
whole bone in Figure \-5. The total length of the
bone is I. If the bone is subjected to a pair of tensile
forces. the length of the -bone may increase to I' or
by an amount .1i\ = I' -I. The normal strain is the
ratio of the amount of elongation to the original
F.,
F,
A
Shear strains are related to distortions caused
shear stresses and arc cornmonly denoted with t
symbol y (gamma). Consider the rectangle (ABC
shown in Figure 1-6 thm is acted on by a pair of ta
gential forces that deform the rectangle into a p
allelogram (AB'C '0). 'If the relative horizontal d
placement of the top and the bOllom of t
rectangle is d and the height of the rectangle is
then the average shear strain is the ratio of d and
which is equal to the tangent of angle y. The angle
is lIsllall~" vcry small. For small angles. the tange
of the angle is approximately equal to the angle
self measured in radians. Therefore, the avera
shear strain is "y = cllh.
Strains arc calculated by dividing two quantit
measured in units of length. For most application
the deformations and consequently the strains
volved may be very small (c,g" 0,001), Strains c
also be gi\'en in percemages (e.g.. O.l%).
STRESS-STRAIN DIAGRAMS
B
F,
Different I11mcrials may demonstrate differe
stress-strain relationships. Consid(~r the stre
strain diagrarn shown in Figure 1-7. There arc
distinct points on the curve, which arc labeled as
P, E, Y, U, and R. Point 0 is the origin of the Sli'e
strain diagram, which corresponds to the initial (
load, no deformation) state. Point P represents t
proportionality limit. Between 0 and P. stress a
strain are linearly proportional and the stre
strain diagram is a straight line. Point E represen
the clastic limit. Point Y is the .\"ield point, and t
stress (T.. corresponding to the yield point is call
the yield slrength of the material. At this Slr
level, considerable elongation (yielding) can occ
without a corresponding increase of load. U is t
highest stress point on the stress-strain diagra
The stress (r is the ultimate strength of the mat
ial. The last point on the stress-strain diagram is
\vhich represents the nq)ture or failure poinl. T
stress at which lhe failure occurs is called the ru
ture strength of the material. For some materials
may not be easy to distinguish the elastic limit a
the yield point. The yield strength of sLieh materi
is determined by the offset method, which is a
plied b.y drawing a line parallel to the linear secti
of the stress-strain diagram that passes through
strain level or approximately 0.2 % • The intersecti
of this line with the stress-strain ClWVC is taken
be the vielel point, and the stress corresponding
this po-int is called the ~\pparent yield strength
the material.
ll
Definition of shear stress. Reprintecl with permission {rom Ozkaya, N, (I 99B). Biomechanics. /11 W.N.
Rom, Environmenlal and Occupational Medicine
(3rd ed" pp. 1437-/454). New York: LippincorrRaven.
I
length, or E = c,11 1. If the length of the bone increases in the direction in which the strain is calculated. then the strain is tensile and positive. If
the length of the bone decreases in the direction
in which the strain is calculated, then the strain is
compressive and negative.
"nfj
,
----,-
I'
F
~
,
"'"
I
--
.; .:.\
~
-
):
-.
I
'.l/
......
F
1
Definition of normal strain. Reprinted with permission from 6zkaya. N. (/998). Biomechanics. In W.N. Rom, Environmental (1nd Occupational !v1edione (lrd ed., pp.
/437-1454). New York: Lippi(lCOfl-R(l~'efl.
•
Note that a given material may behave dilTcr~
ently under different load and environmental con~
ditions. If the curve shown in F'igurc 1~7 represents the stress"strain relationship for a material
under tensile loading, there ma).o' be a similar but
ELASTIC AND PLASTIC DEFORMATIONS
or
Elasticit:-.· is defined as lhe ability
a material to
resume its original (stress-free) size and shape on
removal of applied loads. 1n other words, if a load
different curve representing the stress-strain relationship for the same material under compressive
or shear loading. Also. temperature is known 10411leI' the relationship between stress and strain. For
some materials, the stress-strain relationship may
also depend on the rate at which the load is applied on the material.
u
A
y
I_
'I
d
S'
B
~C
/ - - - - - - - - - 7 C'
/
/
h
/
r1.t /
/
/
/
/
,l-
,
/
AL----;-.-------';0
F
Stress-strain diagrams. Reprinted with permission
Definition of shear strain. Reprinted wirh permission from OZkaya, N. (1998). Biomechanics. In W.N.
Rom, Environmental and Occupational Medicine (3rd
ed.. pp. /437-1454). New York: Lippincou-Raven.
•
from Ozkaya, N. (1998). Biomechanics. In W.N.
Rom, Environmental and Occupational Medicine
(3rd ed.. pp. 1437-1454)., New York: LippincouRaven .
;
is applied on a material such that the Stress generated in the material is equal to or less than th~
elastic limit, the deformations that took place in
the material will be cOlllpletcl.v recovered once the
applied lands arc removed. An elastic material
\vhose stress·strain diagram is a straight line is
called a linearly clastic material. For such a matcthe stress is linearly proportional to strain.
slope of the stress-strain diagram in the e1asregion is called the elastic or Young's rnodulus
of the material. which is commonly denoted by E.
,Therefore, the relationship between stress and
strain for linearly elastic materials is a := E€. This
equation that relates normal stress and strain is
called a material function. For a given material.
different material functions may exist for different
modes or derormation. For example, SOme materials may exhibit linearly elastic belHwior under
shear loading. For such materials, the shear stress
T is linearly proportional to the shear strain y, and
the constant of proportionality is called the shear
modulus, or the modulus of rigidity. If G represents the modulus of rigidity, then ,. = Gy. Combinations of all possible material functions for a
given material form the constitutive equations for
that material.
Plnsticity implies permanent deformations. Materials may undergo plastic deformations follo\ving
elastic deformations when they are loaded beyond
their elastic limits. Consider the stress-strain diagram of a material under tensile loading (Fig.I-7).
Assume that the stresses in the specimen arc
brought to a level greater than the yield strength of
the material. On removal of lhe applied load. lhe
material will recover the elastic deformation that
had taken place by following an unloading path parallel to the initial linearly elastic region. The point
where this path cuts the strain axis is called the
plastic strain. which signifies the extent of perrl1a~
nent (unrecoverable) shape change that has taken
place in the material.
Viscoelasticity is the characteristic of a material
that has both fluid and solid properties. Most materials arc classified as eilher fluid or solid. A solid
material will deform to a ccrLain extent when an
exlernal force is applied. A continuously applied
force on a Ouid body will cause a continuous deformation (also known as flow). Viscosity is n fluid
property thut is a quantitative measure of rcsis·
tance to flow. Viscoelasticity is an example of how
areas in applied mechanics can overlap, because it
ulilizes the principles of both fluid and solid mechanics.
G
E
linearly elastic material behavior. Reprinted wirh
permission from OZkclycl, N. (1998)_ Biomecll<lflics. In
W.N. Rom. Environmental and Occupattonal MecHcme (3rd cd., pp J437-1Li54.J. Ne....,; York: Lippincott-
Raven
VISCOELASTICITY
\·Vhcn they are subjected to relatively low stress
els, many materials such as metals exhibit ela
material behavior. They undergo plastic defor
tions at high stress levels. Elastic materials defo
instantaneously when they are subjected to ex
nally applied loads and resume their original sha
almost instantly when the applied loads are
mo\·cd. For an elastic material, stress is a function
strain only, and the strcss-strain relationship
unique (Fig. 1-8). Elastic materials clo not exh
time-dependent behavior. A different gl'OUp of m
rials, such as polymer plastics, metals at high t
peratures, and almost all biological materials,
hibits gradual deformation and recovery w
subjected to loading: and unloading. Such mater
are called viscoelastic. The response of viscoela
materials is dependent on how quickly (he loa
applied or removed. The extent of deformation
viscoelastic materials undergo is dependent all
rate at which the deformation-causing loads are
plied. The stress-strain relationship for a viscoela
material is not unique but is a f1.lI1ction of time or
rate at which the stresse.s and strains are develo
in the material (Fig. 1·9). The word "viscoelastic
made of two words, Viscosity is a fluid property
,--;
->".
I.'·
"
r
Increasing
suain rale
Ii)
<
Strain
rate~dependent viscoelastic
with an instantaneous strain that would remain at
a constant level until the load is removed (Fig.
I-lOB). At the instant when the load is removecl, the
deformation will instantl)' and completely recover.
To the same constant loading condition, a viscoelastic material will respond with a strain increasing and ciCCI-casing graduall)r. If the material is
viscoelastic solid, the recovery will eventually be
complete (Fig. 1-.1 DC). If the material is viscoelastic
fluid, complete recovery will never be achicved and
there will be a residuc of defOl'mation lerr in the
material (Fig. 1-IOD). As illustrated in Figure
l-11A, a stress·relaxation experiment is conducted
material be-
havior. Reprinted with permission /rom 6zkaya. N.
(1998). Biomedlc1llic5. In WN. Rom, Environmental
and Occupational Medicine (3rd ed., P.o.
1.137-/454). New York: Lippincorr·Raven.
is a measure of resistance to now. Elasticity is" solid
material property. Therefore, viscoelastic materials
possess both nuid- and solid-like properties.
For an elastic material, the energy supplied to
deform the material (strain energy) is stored in the
material as potential energy. This energy is available to return the material to its original (unstressed) size and shape once the applied load is removed. The loading and unloading paths for an
elastic material coincide. indicating no loss of energy. Most elastic materials exhibit plastic behavior
at high stress levels. For e1asto-plastic materials,
some of the strain energy is dissipated as heat during plastic defat-mat ions. For viscoelastic materials,
some or the strain energy is stored in the material
as potential energy and some of it is dissipated as
heat regardless of whether the stress levels are
small or large. Because viscoelastic materials exhibit lime-dependent material behavior. the differences between elastic and viscoelastic material responses are most evident under time-dependent
loading conditions.
Several experimental techniques have been designed to analyze the time-dependent aspects of
material behaviOl: As illustrated in Figure 1-1004, a
creep and recovery test is conducted by applying a
load on the matcl¥ial, maintaining the loael at a constant level for a while, suddenly removing the load,
and obsen;jng the material response. Under a creep
and recovery test. an elastic material will respond
-~
o
Creep and recovery test. Reprinred wirh permission from Ozkay,l, N. (1998). Biomechanics. In W.N.
Rom, Environmental and Occupational Medicine
(3rd ed., pp. 1437-745/1). New York: LippincorrRaven.
,
'by straining the Olalcriallo a level and maintaining
the constant strain while observing the stress response of the material. Under a stress-relaxation
lcst, an elastic mater-ial will respond with a stress
developed insw.ndy and maintained at a consWnl
level (Fig. I-II B). That is, an elastic malcrial will
not exhibit a stress-relaxation behavior. t\ viscoelaslie material. conversely', will respond with an initial
high stress level that will decrease over time. If the
m;terial is a viscoelastic solid, the stress level will
nevcr rcduce to zcro (Fig, I-lie), As illuSlrated in
Fiourc
I-II D, the stress will evct11uallv
o
. reduce to
zero for a viscoelastic nuid.
A
'U
'0
cr
o·u
a
MATERIAL PROPERTIES BASED
ON STRESS-STRAIN DIAGRAMS
PRINCIPAL STRESSES
There are infinitely many possibilities of constructing elements around a given point wilhin a
structure. Among these possibilities, there may be
= E,
to
B
The stress-strain diagrams of two or Il"wrc materials
can be compared to determine \vhich m<:ucrial is rei·
atively stiffer, l1C:lrdcl~ tougher, more ductile, or more
brittle. For example, the slope of the stress-strain di~
agram in the clastic region represents the clastic
modulus that is a measure of the relative stiffness of
materials. The higher the elastic modulus, the stiffer
the material and the higher its resistance to deformation. A ductile material is one that exhibits a large
plastic deformation prior to failure. A britlie material, such as glass, shows a sudden failure (rupture)
without undergoing a considerable plastic deformation. Toughness is a measure of the capacity of a material to sustain permanent defonllation. The toughness of a matedal is measured b~: considering the
total area under its stress-strain diagram. The larger
this area, the tougher the malerial. The ability of a
material to store or absorb energy without permanent deformation is called lhe resilience of the material. The resilience of a material is measured by its
modulus of resilience, which. is equal to the area under the stress-strain curve in the elastic region.
Although thcy arc not directl\' rclated to the
stress-strain diagrams, other important concepls
are used to describe material properties. For cxam~
pie, a material is called homogeneous if its properties do not vary from location to location within the
material. A material is called isotropic if its properlies are independent of direction. A material is
called incompressible if it has a constant denSity.
0"0
C
U:
l110
G
D
10
Stress-relaxation experiment. Reprinted with permission from Ozkaya, N. (1998). Biomechanics. In
WN. Rom. Environmental and Occupational Medicine
(Jrd ed.• P.o. 1437-1454). Ne~··1 York: Lippincott-Raven.
one: element for which the normal stresses
maximum and minimum. These maxin1lrm
minimum normal stresses arc called the princi
stresses, and the planes whose normals are in
directions of the maximum and minimum stJ"e
are called the principal plancs, On a princi
plane, (he normal stress is either maximum
minimum. and the sheal" stress is zero. It is kno
that fracture or material failure occurs along
planes of maximum stresses, and structures m
be designed by taking into consideration the m
/ imulll stresses involved. Failure by yielding
cessive deformation) n.lay occur whenever
largest principal stress is equal to the y
strength of the material or failure by rupture m
occur whenever the largest principal stress ,is
equal to the ultimate strength of the material. For
a given structure and loading condition. the principal stresses ma~" be within the limits of operational safely. However, the structure must also be
checked for critical shearing stress. called the
maximum shear stress. The maximum shear SlI-es$
occurs on a material element for which the normal
stresses are equal.
A
a
I
-
Tension
-
]
o
min
1 cycle
FATIGUE AND ENDURANCE
-
-
(j"
am· - - - ':'.~
(j
Principal and maximum shear stresses are useful in
predicting the response of materials (0 static loading configurations. Loads that Illay not cause the
failure of a structure in a single application may
cause fracture when applied repeatedly. Failure may
occur aher a few or many cycles of loading and unloading, depending on factors such as the amplitude
of the applied load, mechanical properties of the
material, sIze of the structlire, and operational conditions. Fracture resulting from repeated loading is
called fatigue.
Several experimental techniques have been developed to understand the fatIgue behavior of materials. Consider the bar shown in Figure 1-12;:1.
Assume that the bar is made of a material whose
ultimate strength is U'w This bar is first stressed to
a mean stress level (1m and then subjected to a
stress fluctuating over time, sometimes tensile
and other times compressive (Fig. 1-128). The
amplitude (T:, of the stress is such that the bar is
subjected to a maxImum tensile stress less than
the ultimate strength of the material. This reo
versible and periodic stress is applied until the
bar fractures and the number of cycles N to fracture is recorded. This experiment is repeated all
specimens having the same material properties by
applying stresses or varying amplitude. A typical
result of a fatigue test is plotted in Figure 1-12C
on a diagram showing stress amplitude versus
numbct· of cycles to failure. For a given N. the corresponding stress value is called the fatigue
strength of the material at that nun1ber of cycles.
For a given stress level, N represents the fatigue
life of the material. For some matel"ials, the stress
amplitude versus number or cycles curve levels
off. The stress CT, at which the fatigue curve levels
off is called the endurance limit of the material.
Below the endurance limit, the material has a
high probability of not failing in fatigue, regardless of how many cycles of stress are imposed on
the material.
:· · . . .
t
o max - - ..... "7"
-
- .......: ..'
-
-
-
-
-
-
•••••
- ~:.;. - - - ";"~ - - ~:~
••
•••••
••
•••••
lime
Compression
B
°0- - - - - - - - - - - - - L-_,---'---,---,--N
10'
10'
10'
c
Fatigue and endurance. Reprinred with permission
from Olkaya, N. (1998). Biomechanics. In W.N. Rom,
Environmeniat and Occupational Medicine (3rd ed.,
pp. 1437-1454). New York: Lippincou-Raven.
The fatiguc behavior or a material depends on
several factors. The higher the temperature in
which the material is used, thc lower the fatigue
strength. The fatigue behavior is sensitive to surface
imperfections nnd the presence of discontinuities
within the material that can cause stress concentrations. The fatigue failure starts \vith the creation of
a small crack on the surface of the material. which
can propagate under the effect of repeated loads, resulting in [he rupture of [he material.
Orlhopaedic devices lII)dergo repeated loading
and unloading as a result of the activities of the patients and the actions of their 111uscles. Over a pe-
riod of vears. a weight-bearing prosthetic dc\·icc or
fi;.;:ati~n device can be subjected to a consiclerabk
number of cycles of stress reversals as a result or
noHnal daily activity. This cyclic loading and un~~ .•l;"N can cause faLigue failure of the device.
a'
b~"
clinicians (Q provide an introducLor~: level
knowh:dge about each joint s~'stem.
PART III: APPLIED BIOMECHANICS
A new section in the third edition of this book
Biomechanics
the Musculoskeletal System
even a simple task c.'\ecuted b.v the
musculoskeletal svstcm requires a broad. in-depth
:~;' knowledge of various fields that ma~' include 1110tor control, neurophysiology, physiology. physics.
and biomechanics. For example, based on the purpose ancl intention or a task and the sensOl'~' information gathered from the ph~'sical cndronmcnl
and orielllatioll of tilL' body and joints, the central
nervous system plans a strategy for a task execu:;.: lion. According to the strategy adoptc:d. Illuscles
/ . '.,' will be recruited lO provide Lh<..' forces and 1110mcnts required for the movement and I.Jalance of
the s.)'slem. Consequently, the internal forces will
be changed and soft tissues will experience different load conditions.
The purpose
this book is to present a \\'cll~
balanced synthesis of information gatllCred frorn
various disciplines. pro\'iding a basic understanding
of biomechanics of the musculoskeletal system. The
material presented here is organized (0 cover three
areas of musculoskeletal biomechanics.
;~;
or
PART I: BIOMECHANICS OF TISSUES
AND STRUCTURES
The material presented throughout this textbook
provides an introduction to basic biolllechanics of
the musculoskeletal system. Part I includes chapters on the biomechanics of bone. articular cartilage, tendons and ligaments, periphcral nerves, and
skeletal muscle. These are augmcnted wilh case
studies to illustrate the imponarll concepts for understanding the biomechanics of biological tissues.
PART II: BIOMECHANICS OF JOINTS
Part II of this textbook covers the major joiots of the
human body, from the spine to the ankle. Each
chapter contains information about the structure
and functioning of the joint. along with case studies
illuminating the clinical di~lgnosis and management
of joint injlll)' and illness. The chnpters are written
troduces important issues in applied biomechani
These include the biomechanics of fracture fixati
arlhroplasty; sitting, standing. and lying; and gait
is important for the beginning studenl to und
stand the application or biomechanical principles
clirfcrcnt clinical areas.
Summarv
1 Biomechanics is a young and dynamic fidd
study based on the recognition thaI conventio
cllginccl'ing thcorks and methods can be useful
understanding and solving problems in physiolo
and medicine. Biomcclwnics considers the appli
tions of classical rncchanics to biological problem
The flcld of biomechanil:s flourishes from the co
eration among life scientists, physicians, enginee
and basic scientists. Such cooperation require
certain amount of common vocabulary: an engin
must learn some anatom~: and ph)'siology, nnclm
ical personnel need to understand some basic c
cepts of physics and mathematics.
2 The information presented throughout t
textbook is drawn from a large scholarship. The
thors aim to introduce some of the basic concepts
biolllechanics related to biological tissues a
joints. The book does nOl intend to provide a co
prehensive review of the literature, and readers
encouraged to consult the list of suggcsted read
below to supplcmcnt theil' knowledge. Some ba
textbooks arc listed here, and studcnts should c
sult peer-revicwed journals for in-depth presen
Lions of the latest research in specialty arcas,
SUGGESTED READING
Black. J. (19SS). Onhop'lcdic Bionmleriuls in R,'sl,.'ardl and P
til:c. New York: Churchill Li\·in!!stonc.
Brollzino. J.D. (Ed.) (1995). The -Biom(.'dic<ll Engincl.'ring H.
book. 80(:01 R:llOn. rL: CRC Press.
B(lrst('in, A.H., & \Vright. T.~'I.( 1993). Fundamellt;lls of Or1hop<
Biolllc:dmnics. Ballirnorl': Williams & Wilkins.
Chaffin. 0.8.. & All{krsson. G.B.J. (1991). O<::cupational Bio
ch~lllics (2nd l'<'I.). Nl'\\' York: John Wiley & Sons.
Fung. Y.c. (1981). Biolllec!wnics: Mechanical Properties of Li
Tissul.·s. New York: Springl.'r·\\:rJag.
FUll!!.
yc. (1990). Biomcch~l1lics: ;\·Iotioll, Flo\\" Slrcss, :llld Gro
New York: Springl'r,Vl'rlOIg.
... :':.
H~l\'. J.G ..
6:.. RI..'id. J.G. (1988). .'\nalOllw. '\lcchanil:s .md Human Mo-lioll (2nd cd.). Englewood Cliffs, NJ: Pn:llIkc-Hall.
Kelly, O.L. (1971). Kinl.'siology: FLllldaml.'llWls of \Iolion D~·StTip·
lion. En!.:.lcw()od Cliffs, NJ: Prelltice-Hall.
\Iow,
Vc.. &:
Hayes,
w.e.
(1997). Basic Orthopaedic Biorncch.mics
(2nd cd.). New York: Raven Press.
,\low. \I.e., Ratdiff, A".& Woo. S.L.·Y. (Eds.). (1990). BicHlll.:chanics
or {)i~tl"(llrodial Joinls. Nr.:\\' York: Sp.-jnga-Verlag.
NahulIl. A.M .. &. Md\'in. J. (Etls.). (19S5). The BiOlllcch:lI1il:s of
TI'UUll'l.
Norwalk. CT: l\ppl('ton-Ct:llhll~'.Crofts.
Nordin, M .. Andersson. G.B.J., & Po\X", M.H. (Eds.). (1997). ~plllSCll­
1()sk..:k'IJI Disonk'J1; in the \VorkpbcL'. Philadclphin: :\'1osby,Ycilr
Book.
Nordin. ~L & Franb,:1. \l.H. (Eds.l. (1989). Basit- OiolUl.'chanics of
lhe :\hlst.:uloskdctal S,\':o;'lClll (2nd t:d,), Philaddphi:l: 1...:<.\ &
F.:bi"'.:r
OzkaYil,eN.·( 1998). Bi(Hn~d1iJnics. III W.N. Rom, EnvirorH1H:nlal and
OCCllp<:lliollal :\kdicill~' Ord cd., pp. 1437-1454). New York:
Li ppi m:olt -RaVCll.
OZk:IY;\. N.. & Nordin. M. (1999). Fundamelltals of Biolll<.'ch'lllics:
Equilihrium, :\lotioll. <lnd Dcfonlwtioll (2nd I..'d,). Nt:\\" York
Sprin!.!(,I~Vl.'rlag.
Schmid.Schonbl'i;l, G.\\'., \\'00. 5.1...·).... & Z\\"eifach. B.\V. (Eds.).
(1985). Frontiers in Biomechanics. ~·h.'\\" York: Springer.Verlag:,
Skal<:lk, R" & Chkn, S. (Eds.). (!98i). Hllndbonk or BiOi..'ngincering.
New York: McGraw·Hill.
Thompsoll. C. W. (1989). :\·tanu'll of Stlllt'lur<ll
cd,). 51. Louis. MO: ·fill1l..'s Mirror/:\·Iosbv.
Kinl~si()log~'
(11th
Williams. M., & lissner, H.R. (1992). Bion;l.'chanics of !·luman ;\'tolion (3r<l cd.). Phibddphin: Sallll(k:rs,
Wintcl'. D.!\. (1990). Biol1lcchnnics :Illd ~:lolor Control or Hum"n
Behavior (2nd I..'d.). New York: John Wiky &: Sons.
Willlel'S. J,!lil.. &. Woo. S.L-Y. (Eds.), (1990). Multiple Muscle SysIl'IHS. New York: Springcr.verktg.
The System
International
d'Unites (51)
Dennis R. Carter
The 51 Metric System
Base Units
Supplementary Units
Derived UnitS
Specially Named Units
Standard Units Named for Scientists
Converting to 51 From Other Units of Measurement
fined in terms of the \V~\'c1cngth of radiation emitted from the krvpton-S6 atom.
The''S! Metric System
,,;The System Intemational d'Unites (SI), the metric
system, has evolved into the most exacting system of
'measures devised. In this section. the 51 units of
lllcasurcmcnl used in the science of mechanics arc
described. SI units used in electrical and light sciences have been omitted for the sake or simplicity.
~__
• fifo'
<""'i,ilY ' BASE UNITS
--~
The 51 units can be considered in three groups: I,
tbe base units; 2. the supplementary unils; and 3.
'j;", the derived units (Fig. App-I). The base units are a
,', small group or standard measurements lhal have
been arbitrarilv defined. The base unit for length is
'-'rhe meler (Ill), "and the base Llnit for Illass is the kilogram (kg). The base units for time and temperature
are the second (s) and the kelvin (Kl. respectively.
Definitions 01" the base units have become.:: increas·
.ingly sophisticated in response to the expanding
needs Hnd capabilities of the scientific community
(Table App-I). For example, the meter is now de-
SUPPLEMENTARY UNITS
The radian (rad) is a supplemental)' unit (() measure
plane angles. This unit, like the base units, is arbitrarily defined (Table App-I). Although the radian is
the 51 unit for plane angle. the unit of the degree
has been retained for general use because il is firmly
established and is widely Ltscd "round the world. A
degree is equinllcnl to rrll80 rad.
A"
QUANTITIES EXPRESSED
IN TERMS OF UNITS FROM
WHICH THEY WERE DERIVED
DERIVED UNITS
ivlostunits of the 51 system are derived unils, meaning lhat they- are established from the base units in
accordance with fundamcntnl physical principles.
Some of these units are cxp['css~d in terms of the
base units from which they are derived. Examples
aloe area, speed, and acceleration. which arc expressed in the Sf units
square meters (m~). meters
per second (m/s), and meters pCI' second squared
(In/s 2), respectively.
or
DERIVED UNITS
WITH SPECIAL NAMES
FORCE
new Ion
kg m/s 2
MOMENT
OF FOACE
ACCELERATION
speED
DENSITY
VOLUME
AREA
qp
,- - . .
PRESSURE & STRESS
pascal
N/m 2
~\
. \l/~
\' Il/' _~r;)
~~ ~,I //'/~'"
~
;~:::::<':~:~~J" J
,
,~/~,:'!j
~DERIVED~
,~m'?I~~'TS
i~
ENERGY & WORK
joule
Nm
POWER
watt
JIS
TEMPERATURE
degree Celsius
K'- 273.15
radian (fad)
PLANE ANGLE
----SUPPLEMENTARY UNIT
meter (m)
LENGTH
kilogram (kg)
MASS
second (s)
TIME
kelvin (K)
TEMPERATURE
BASE UNITS
The International System of Units.
19
The 51 unit of pressure, the pascal, is therefore
defined in terms of the base 51 units as:
Specially Nanzeel Units
Other dedved units are similarl.v established from
the base units but have been given special names
(Fig App-I and Table App-I). These units are defined
through the lise of fundarnental equations or physical laws in conjunction with the arbitrarily defined
Sf base units. For example, Newton's second law of
motion states that when a body that is free to Il10vC
is subjected to a force. it will experience ~m ~lCcelcr­
alion proportional to thai force and inversely proportional to its own mass. i\Jlathcmatically, this principle can be expressed as:
force = Illass X acceleration
The Sf unit of force, the newton (Nl, is lherefore
defined in terms of the base SI units as:
1N
=
1 kg X I
I11/S:!
The Sf unit or pressure and stress is the pascal
(Pa). Pressure is defined in hydroslaties as the force
divided by the area of force application. Mathem4llicall~l, this can be expressed as:
pressure = force/area
~._.m
..__ .._
I Pa
= IN I I
111"
Allhough the S[ base llnil of temperature is the
kc.:lvin, the derived unit of degree Celsius (OC 01' c) is
much marc commonly used. The degree Celsius is
equivalent to the kelvin in magnitude. but the ab~
solute value of the Celsius scale differs frol11 that o
the Kelvin scale such that °C = K - 273.15.
,",Vhen the 51 s~'stcm is used in a wide variely o
measurements, the quantities expressed in terms
of the base. supplemental. or derived units ma~t be
either very large or very small. For example, the
arca on the head of a pin is an extremely small
number when expressed in terms of square meters
Conversely, the weight of a whale is an extremely
large number when expressed in terms of newtons.
To accomrnodate the convenient representation o
small or large quantities, a system of prefixes has
been incorporated into the SI system (Table
App-2), Each prefix has a fixed meaning and can
be used with all 5Iunils. \!\Then used with the name
or the unit, the prefb: indicates that the quanti!}!
described is being expressed in some multiple o
__.__._.__.. _
I
; Definitions of Sl Units
Base 51 Units
meter (01)
The meIer is the length eqllal to 1,650,763.73 wavelengths in vacuum of the radiation
corresponding to the transition belween the levels 2PHi and Sd" of the krypton-86
atom.
kilogram (kg)
The kilogram is Ihe unit of mass and is equallO the mass of the international proto·
type of the kilogram.
second (s)
The second is the duration of 9.192,631,770 periods of Ihe radiation corresponding to
the transition between the two hyperfine levels of the ground state of the cesium-133
atom.
kelvin (k)
The kelvin. a unit of thermodynamic temperature, is the fraction 1/273.16 of the thermodynamic temperature of the triple point of water.
Supplementary SI Unit
radian (rad)
The radian is the plane angle between two radii of a circle that sub tend on the circum~
ference of an arc equal in length to the radills.
Derived SI Units With Special Names
newton (N)
The newton is that force \tvhich, when applied to a rpass of 1 kilogram, gives it an aCceleration of 1 meter per second squared. 1 N= 1 kg rn/s / .
pascal (Pa)
The pascal is the pressure produced by a force of 1 newton applied. with uniform distribution, over an area of 1 square meter. 1 Pa = 1 N/m~.
joule (J)
The joule is the work done when the point of application of a force of 1 newton is
displaced through a dist~nce of 1 meter in the direction of the force. 1 J = 1 Nm.
wall (W)
The watt is the power that in 1 second gives rise to the energy of 1 joule. 1 W ::: 1 J/s.
degree Celsius (C)
The degree Celsius is a unit of thermodynamic temperature and is equivalent to K -
273.15.
Standard Units Nmned
for Scientists
Factors and Prefixes
SI Prefix
SI Symbol
giga
G
mega
Iv1
kilo
k
hecla
h
deka
da
deci
d
centi
c
rniJli
rn
micro
I'
nana
n
pica
p
One of the more interesting aspects of the SI
is its lise of the names of famous scientists a
dard units. In each case, the lInit was named
scientist in recognition of his contribution
Geld in which that unit plays a major role
Reprinred wirh permission from Ozkaya. N.. & Nord,-n. M.
(1999). Fundamentals oi Blomechani(!>: Equilibrium. Mo·
tion. and Deiormatlon (2nd ed.) New York: Springer·lJer/ag.
p, iO.
ten times the unit used. For example. the millime·
'tel' (mm) is used to represent one thousandth (10"·1)
of a meter and a gigapascal (Gpa) is uscd to denotc
one billion (10') pascals.
App-3 lists a number of Slullits and the scien
which each was named.
For example. the unit of force. the neWlo
named in honor of the English scientist Si
Ncwlon (1624-1717). Hc wa' cducalcd 'II
College at Cambridge and later rclurned to
College as a professor or mathematics. Early
career, Newton made fundamental contribut
malhcmalics Ihat formcd Ihc basis of diffc
and integral calculus. His other major disc
were in the fields of optics. astronomy. grav
and mechanics. His work in gravitation wa
portedly spurred by being hit on the head by
ple falling from a tree. It is perhaps poetic
that the SI unit of one newton is approxi
equivalent to the weight of a medium-sized
Newton W~lS knighted in t 705 by Qucen i\lt
his monumental contributions to science.
---_._--------
, SI Units Named After Scientists
Symbol
Unit
Quantity
Scientist
Country of Birth
Oates
A
1775-1
1736-1
1701-1
1791-1
1797-1
1857-1
1818-1
1824-1
1642-1
1787-1
1623-1
1823-1
1856-1
1745-1
1736
1804-1
ampere
electric current
Amphere, Andre·tvtarie
France
(
coulomb
electric charge
Coulomb, Charles Augustin de
France
O(
degree celsius
temperature
(elsius. Anders
Sweden
F
farad
electric capacity
Faraday, Michael
England
H
Hz
henry
inductive resistance
Henry, Joseph
United States
hertz
frequency
Hertz, Heinrich Rudolph
Germany
J
joule
energy
Joule. James Prescott
England
K
kelvin
temperature
Thomson, William (lord Kelvin)
England
N
newton
force
Newton, Sir Isaac
England
fl
ohm
eleclrlc resistance
Ohm. Georg Simon
Germany
Pa
pascal
pressure/stress
Pascal, Blaise
France
5
siemens
electric conductance
Siemens, Karl Wilhelm (Sir William)
Germany (England)
T
testa
magnetic flux density
Testa. Nikola
(roatia (US)
volt
electrical potential
Volta, (ount Alessandro
Italy
walt
power
Watt, James
Scotland
weber
magnetic flux
Weber, Wilhelm Eduard
Germany
V
W
Wb
Conversion of Units
Moment (Torque)
Length
1 centimeter (em)
=
0.01 meter (01)
1 inch (in) = 0.0254 rn
1 foot
(ft)
= 0.3048 rn
1 dyn-cm
=
i Ibf-ft
1.356 N-m
=
i 0. 7 N-rn
Work and Energy
s'
1 yard (ydl = 0.9144 rn
1 kg-m' /
1 mile = 1609 m
1 dyn-cm = 1 erg = lO-} J
1 angstrom (A) = 10" rn
1 Ibi-It
Time
= 1 N-m = 1 Joule IJ)
= 1.356
J
Power
=
I minute (min) :::; 60 second (s)
I kg· rn' / s' = 1 J/s
1 hour (h) = 3600 s
I horsepower (hpl = 550 Ibl·IVs = 7461"1
1 day (dl = 86400 s
1 Watt (1"1)
Plane Angle
Mass
J degree ("') .::;: 1</180 radian (fad)
1 pound mass (1 brn) = 0.4536 kilogr~lm (kg)
1 revolution (rev)
1 slug ::: 14.59 kg
1 rev;::: 211: raci -."-' 5.283 rad
:=:
360"
Temperature
Force
1 kilogram force (kgf) = 9.807 Nevvton (f\J)
<>( '" "r~ ~
273.2
1 pound iorce IIbO = 4448 N
I dyne (elyn) " I0
~I
Pressure and Stress
1 kgl rn·s:' = 1 N/m" = 1 Pascal (Pal
I
Ibi / in' (psi) = 6896 Pa
I Ibl / ft' (psO = 92966 Pa
I dyn / em' = O. I Pa
Reprinted with pefFnisslon from OZkaya. 1'1., & No~dJn. l'1i. (1999). Fundamentals oi Biomechanics: EqUllib·
rium. Motion. and DeformatIon (2nd ed.) New York: Spnnger-verlag. p. 11.
The unit of pressure and stress. the pascal. was
named after the French physicist, mathematician,
and philosopher Blaise Pascal (1623-1662). Pascal
conducted important investigations on the <.:harncteristics of vacuums and barometers and also invented a machine that would make mathematical
calculations. His work in the area of hydrostatics
helped lay the foundation for the later developmenl
of these scientific ficlds_ In addition to his scicntific
pursuits, Pascal was passionalely interested in religion and philosophy and tllllS wrote extensively on
a wide range of subjects.
The base unit of temperature, the kelvin. was
named in honor of Lord Vv'illiam Thomson Kelvin
(1824-1907). Named William Thomson, he was cd·
ucated at the Universit.\· of Glasgo\\" and at C
bridge University. Early in his career, Thol11son
vcstigated the thermal propcnies of steam at a
entific laboratory in Paris. At thc age of 32,
returned to Glasgo\\" 10 accept the chair of Nalu
Philosophy. His meeting with James JOllie in 1
stimulated interesting discussions on the natur
heat, which eventually lecl to the establishmen
Thomson's absolute scale of temperature, the Ke
scale. In recognition of Thomson's contribution
the field of thermodynamics, King Edward VIl c
fen-cd on him the title of Lord Kelvin.
The commonly llsed unit of temperature, the
gree Celsius, \\ as named aftcr the Swedish
lronomcl and lmenlO! Anders CelsiLl' (1701-17
Celsius was appointed professor or astronomy at the
Inii"p-",i,\' of Uppsala at the age of 29 and remained
the university until his death 14 years latec [n
1742, he described the centigrade thermometcl- in a
paper prepared for the Swedish Academ!' of SciThe name of the centigrade temperature
waS officially changed to Celsius in 1948.
to Sf From Other
of'Measurement
Box App·t contains the formulae for the conversion
~f measurements expressed in English and non-Sl
units into Sf units. One fundamental source of
connJSlon in converting from one system to another
is that (wo basic t~'Pes of Illeasun:menl systems exist.
In the "ph\'sical" system (such as SI). the units of
length. time, and nUlSS arc arbitrarily defined, and
other units (including force) are derived fron1 these
,. base units. In "technical" or "gravitational" systems
(slich as the English system). the units of len
time, and force arc arbitrarily' defined, and o
units (including mass) are derived from these b
units. Because lhe units of force in gravitational
tems are in fact the It'eights of stan(~"\rd masses.
version to 51 is dependent on the acceleration
mass due to the Earth's gravity, By' internati
agreement. the acceleration due to gravity
9.806650 m/s'. This \'aille has been lIseel in estab
ing some of the conversion factors in Box App-I.
REFERENCES
F~ircr,
J.L. (19i7). SI .\leu;£' /-land/wok. Nt'w York: Ch
StTibll('r'S Sons.
Ozbva, N., ~ Nordin, ~'l. (1999). FWlfl(llllell1o!s
or
Bi
c/;(wics: EquilihriulJI, .\lotioll, (lut! DejtmJ/(//ioll (2nd
N<:w York: 5pringcr-Vl·r1ag.
Pennychuick, C.J. {1974l. lIulldy .\!tllrif:l's oj" VI/il Crntl't'
Pi/oon (or Riolof!..\' amI .\lcclulJI;cs. New York: John W
&:
Son~.
\Vorld Health
Org~lnizalion.
jt's$iol1s. Gl.'llt'\"C WHO.
(19i7). The SI
/01' t!le
fha/til
Biomechanics
of Tissue and
Structures of the
Musculoskeletal
System
Biomechanics of
Bone
Victor H. Frankel, Margareta Nordin
Introduction
Bone Composition and Structure
Biomechanical Properties of Bone
Biomechanical Behavior of Bone
Bone Behavior Under Various loading Modes
Tension
Compression
Shear
Bending
Torsion
Combined l.oading
Influence of Muscle Activity on Stress Distribution in Bone
Strain Rate Dependency in Bone
Fatigue of Bone Under RepetitivE: Loading
!nfluence of Bone Geometry on Biomechanical Behavior
Bone Remodeling
Degenerative Changes in Bone Associated With Aging
Summary
References
Flow Charts
l
_
""''''''----------.--.:~.",-----
lJitijoduction
TbJ:i~~rpose or the skeletal system is to protect inte.r~.~V:"organs, provide rigid kinematic links and
~~ls:ctEtattachmcnt sites, and facilitate Illuscle ac·
:·:::;:~"pJ:tf91iYli~9,body movement .. Bone has unique stl·UC·
'-:/:::"'{:<'Y~lit~I.-.rI1d mechanical properties that allow it to
,,:~,<,:;;::,/,):',si;(rrY()llt;these roles. Bone is among the body's
,::)},:;;ll~lrci~~tstrllctllres; onl)' dentin and enamel in the
:,:;:::';:,:::':i:,..'./t~¢th.>qreharder.
It is one of the most dynamic and
::",';;::;.~/'i,1;i',.~t. ?9Pp'.cally active tissues in the body and re·
>.";'\1.lAiri,S::~\ctiv0 throughollt life. A highl~; vascular tiss~e'.-.it has an excellent capacity for self-repair and
~lter its properties and configuration in rc. spOI~~e" to changes in mechanical demand. For example, changes in bone dcnsit~1 arc commonly
obsei~·ed after pL:riods of disuse and of greatly increased use; changes in bone shapL' are noted dur·
iog fracture healing and after certain operations.
.;: Thus, bone adapts Lo the mechanical demands
placed on i l.
This chapter (iL'scribcs the composition and
structure of bone tissue, the mechanical properties
of bone, and the behavior or bone under different
loading conditions. Various factors that affect the
mechanical behavior of bone in vitro and in vivo
also are disclissed.
tan
·-'1
Bone Composition and Structure
Bone tissue is a specialized connective tissue whose
solid composition suits it for its supportive and proteclive roles. Like other connective tissues, it con·
sisrs of cells and an organic extraccllular matrix of
fibers and ground substance produced by the cells.
The distinguishing reature of bone is its high con·
tent of inorganic materials, ill the form of mineral
salts, that combine intimately with the organic matrix (Buckwalter et al., 1995). The inorganic compo·
nent of bone makes the tissue hard and rigid, while
the organic component giv~s bone its flexil)ility and
resilience. The composition of bone dirrers depending on site, animal age. dietar:.y histol)', and the pres·
ence or disease (Kaplan et aI., 1993).
In normal human bone, the mineral or inorganic
portion of bone consists primarily of calcium and
phosphate, mainly in the form of small crystals rc·
sembling synthetic hydroxyapatite crystals \vith
the composition Ca",(PO,)o(OI-l),. These minerals,
which a~count for 60 to 70% of ilS dry weight, give
bone its solid consistency. "Vatcr nccounts for 5 to
SOk and the organic matrix makes tip the J'cmainder
.. , ..~l.'
of the tissue. Bone scn·cs <.lS a rcsen/Oir for esse
minerals in the bod~·. particularly calcium.
Bone minend is embedded in variousl~« orie
fibers of the protein collagen, the fibrous po
of the extracellular matrix-the inorganic ma
Collagen ribers (type l) are tough and pliable
they resist strelching ~lnd have lillie L'xtensib
CollagL'1l composes ~\PPJ'().\.irnatdy 9(YYo of the
tracellular matrix and accounts for approxima
25 lO 3W>'r' of the dr~! wl'ight of bone. A univ
building block of the body, collagen also is
chie!" fibrous component of other skddal s
tures. (A detailed descriplion 01" the microstnlc
and mechanical behavior of collagen is provide
Chapters 3 and 4.)
The gelatinous ground substance slIlToun
thc mineralized collagcn fibers consists mainl
protein polysaccharides, or gl~"cosarninogl
(GAGs), primarily in the form of cOlllplt:x ma
molecules called protcoglycans (PGs). The G
s~rvc as a cementing substance between laye
mineralized collagen fibers. Thcs~ GAGs, a
wit.h various noncollagcI1ous glycoprotcins, co
tute approximately 5% of the extracellular rna
(The structure or PG:-:, whi,:h arc vital compon
of artie, i:il' '_·;l~·{iL\gt·. is described in delail in C
kr .) ..1
\Vater is fairl~: abundant in live bone. accoun
for up to 25% of its total weight. Approxim
85°10 of the walt.'r is found in the organic ma
around the collagen fibers and ground sllbsta
and in the h.vdl·ation shells surrounding the
crysl~ds. The other 15% is localed in the canals
cavities that house bone cells and carry nutrien
t h~ bone tissuc.
At the microscopic level, the fllndamcmal S
lllralunit of bone is the osteon, or haversian sy
(Fig. 2·1). At the center of each osteon is a s
channel, called a haversian canal, lIlat con
blood vessels and nerVe fibers. The osteon itself
sists of a concentric series of layers (lamella
mineralized rnatrLx surrounding the central can
configuration similar to growth ring~ in a
trunk.
Along the bOllndflries of each layel: or lamella
small cavities known as lacunae, each cOl1la
011(.' bone cell. or ostcocyte (Fig. 2-1 C). Nume
small channels, called canaliculi, radiate from
lacuna, connecling the lacunae or adjacent lam
and ultimately reaching the haversian canal.
processes extend from the osteocytes into the ca
culi, allowing nutrients '[Tom the blood vessels i
haversian canal to reach the osteocyles.
Lamellae
OsteocYle~
Lacuna-----
B
A
A. The fine structure of bone is illustrated schematically
in a section of the shaft of a long bone depicted without inner marrow. The osteom. or haversian systems.
are apparent as the structural units of bone. The haversian canals are in the center of the osteons, which form
the main branches of the circulatory network in bone.
Each osteon is bounded by a cement line. One osteon is
shown extending from the bone (20x). Adapted from
Basset!, CAL. (1965). £/ecrrical effects in bone. SCI Am,
213.18. B, Each osteon consists of lamellae. concentric
haversian canal. Adapled irom Torrora G.J.. & Anagno
takas. N.P. (198:1). Principles of Anatomy and Physiolog
edJ. Ne~·v York: Harper gRow. C, Along the boundar
the lamellae are small cavities known as lacunae, e
of which contains a single bone cell, or osteocyte.
ating from the lacunae are tiny canals, or canalicu
into which the cytoplasmic processes of the osteoc
extend. Adapted from Torrora G.;.. & AOclgflosrak05. N
(1984). Principles of AnalOmy and Physiology (4th edJ.
York: Harper & Ro~"/.
rings composed of a mineral matrix surrounding the
•
At the pcriphcl)! of each osteon is a cement line,
a nrtrrow area of cement-like ground SubSlrtllCe
composed primarily of GAGs. The canaliculi of Ihc
osteon do not pass this cement line. Like the canaliculi, the collagen fibers in the bOl1e matrix interconnect from one larndla to another within an osteon
but do not cross the cement line. This intertwining
of collagen fibers within the osteon undoubtedly in-,
creases the bone's resistance to mechanical stress
and probably explains _~vhy the cement line is the
weakest portion of the bonc's microstructure.
A typical osteon is approximately 200 micr
tcrs (J..l) in (lin·rnctcr. Hence, evclY point in th
teon is no more than 100 J..llll from the central
catcd blood supply. In Ihe long bones, the os
usually run longitudinally. but they branch
quelltly and anaslOmose extensively with
othcl:
Jnterstitial lamellac span the regions bet
complete O!"h:ons (Fig: 2-1..1). They arc contin
with the ostcons and consist of the same mater
a difTerent geol11ctric configuralion. As in th
!
Frontal longitudinal section through the head. neck,
greater trochanter, and proximal shaft of an adult femur. Cancellous bone, with its trabeculae oriented in
a lattice, lies within the shell of cortical bone.
Reprinted with permission from Gray, H. (T 985). Anatomy
of the Human Body. (73(h American ed.J. Philadelphia:
Lea & Febiger.
1
teons, no point in the interstitiallamellae is farther
than 100 JLm from its blood supply. The interfaces
between these lamellae contain an alTa)' of lacunae
in which oSleocytes lie and from which canaliculi
cXlcnd.
!\l the macroscopic level, all Qonc,~ ..,!re.. c.Qrnposed
of twotvpes, ,00~os'se-qus,~isslle: cortical. or compact,
bone ~\n~d ci\'nc-c-il~·~IS.-Ol:"tl'abeclllal~"'iJ()nc'(FE~-"-i~·2).
COl'tiC,,1 hoi1"Tili'ms the Otlt",:,11-"II:or' cortex: of the
b()I"1~,tll~Cl.. has a ~l,~nse stt'llct,llre ,Si..11]i,ic.\I:t~)th~\i of
ivory. C;ln~~Tlou's bo;;e within ihi~~I;ell is composed
or-thin plates. o'j"lrabeclllae, in a loose mesll.-,~t1·.uc­
lUI~e; lhf([rlier'si-i·ccs--bctwe.en the lrabct:ulat:: are filled
\\~hh red ;;all'O\';" (Fig. 2~j). C~I~~~ii~ll~-b~-~~~~e
isaiTangeCi
A, Reflected-light photomicrograph of cortical
from a human tibia (40;<). B, Scanning electron p
tomicrograph of cancellous bone from a human tibi
(30)-:). Reprinted wirh permission from Carrer. D.R., & Na
We. (1977). Compact bone fatigue damage. A microsco
aminalion. Clin Onhop, 127, 265.
in concentric lacunae-cQn'i"ainTilgJii!l)el-
lae--ouiclocsn-ot em~l;-in haversian canals. The osteoc:\;tes re'ceive nUlricn-islhrough canaliculi from
blood vessels passing through the red marrow. ~.Q.r­
lieal..bone always SUITOt.II).~~~ G!.pccl!.ous lJ.one. bUl·
thi-rdative quantily of each lype varies' among
bones and wilhin individual bones according to
funclional requirements.
On a rnicroscopic level, bone consists of .w
and lamellar-bone (Fig. 2-4). Woven bOlle is-co
erecl"immatllre bone. This type or bone is rOll t1 d
enlbryo.-in the newborn, in the- fracturc'callus,
lhe riietaphysial region of growing bone as wel
tUIl-lOrS, ostcogcl]csis lmperfecla, and pagctic
Lal-nellar bone begins to form 1 month arter bir
activelv replaces woven bon~. Lamellar bone is therC'for~ a "marc" l11all~'c bOlle.--All bones are §urrounded by a dC.llse __ U_bJ~olls
membrane called the periosteulll (Fig. 2-1.-\). Its
ouler fa)~ei: (~ -pe-i::nl-eat-e-(Ch~);-'-GTood vessels (Fig.
2-5) and nerve fibers that pass into the cortex via
Volk,llunn's cH"nalS: c()r\-i1c-EUn·g~\\;i.t~-1 thc--rl~i\.;e,:sian
eLl m\ls '\Il'(Cc',~,~~,6~,lir~~Lt(? tll~-,s,a.·!1·~,~,1..1('-li~ ,I)one .. f\ n
i n I'~<;-l~" o~_~,~~?g,~n .i..~,J,_~ver C()ll,t,,~1.'i Ils '" l?5)Il~ ~t:llsIY~
spo"nsible.....",'(or ,5:_".".
~(;nen~ting
nc\v"" •.Gc;'n-c'''du'i-ing
gro\Vtl~,
. .",..
'-""'-"""~"""""""
-"""
.
-
which m":Cco\;'cr'ccT'witTl arlicufm- cartilage. I
long borles, 'a- thin'llcr "nl'e'moi~a-ne:llle-cndo
v:,\'
line";; t heee n 1 ral (;;;";;clujIaty j ~,,",;il
Ii icll;~
\VilTl'-y~cfl~~~~;~ fall'; -iiuirio\\':~ The' endostcum
tai,'ls~~~s:l'~~;~l;ia~t; ,~,~d .alse),. giant n~-~,fli"'l-ll1-c'
l.ioli"-e"lisctlTfec! osteoela-sfs: both of which
i m pOI:f~li1i'I~(;fc's"'Ttl-dlc'l~cm odc ling aMi1~ir"(:cso I'
of bone .
-
- -•
•
-
and ,'cRail'. (ostcoblasls). The periosteum c
th~cnli~ -b0I1::' except for the - joint-'Slld"
-
'--
-
"
-
Lamellar
~ .f v~~, J
/'i .\ } ,
l
,~. \. \ \)
'("'1::.;,
Woven
Schematic drawing and photomicrographs of lamellar and woven bone, Adapted
from Kaplan, F.5., Hayes, We., Keaveny. T.M., er a/. (1994). Form and funcrion of bone.
In S.R. Simon (Ed.). Orthopaedic Basic Science (pp. 129, 130). Rosemonr, fL: AAOS.
.~
Photomicrograph showing the vasculature of cortical bone. Adap,ed from K~lplan,
w,e.. Keaven}~ rA1 .. et M (1994). Form and function of bone. In 5.R, Simon
(Ed.). Orthopaedic Basic Science (p. /3i}. Rosemon£, It: AAOS.
FS .. Nayes,
•
Biomechanical Properties
of Bone
Biomechanically, bone tissue may be regUl~ded a~ a
t\v9~ph~,lse (biphasic) c01JlposiL(;~ I1),aterial;" with the
mineral as onCjJhase and the collagen and ground
substance as_.the o~her. In such matcl"'ials (n nonbio~­
logic-al exal,'lple is fiberglass) in which a strong. brittle material is embedded in a w~akcc more Oexible
one, the combined substances ;:tre stronger for their
weight than is either substance alone (Bassett,
1965),
Functionally, the most important mechanical
properties of bone are its strength and stiffness.
These and other characteristics can best be understood for bone, or any other structure, by e:'\amining its behavior under .loading, that is, under the innuence of externally applied forces. Loading causes
a deformation, or a change in the dimensions, of
the~-strllctllre. \·Vhen a load in a known direction is
imposed on a structure, the deformation 01" that
structure can be rncasurcd and ploued 011 a loaddeformation curve. I\lluch information about the
strength, stiflness, and other mechanical proper
of the structure can be gained b~' examining
curve.
A hypothetical load-deformation curve for
somewhat pliable fibrous structure, such as a l
bone, is shown in Figure 2·6, The i1~~lial (strai
line) portion of the curve, the c1asticJ'cgion. reve
the elasticity bf the structure, th:u is, its capacity
returning io its original shape after the load is
moved, As· the load is applied, deformati911 occ
btll is not permanent; th-e structure recovers its o
inal sllape \\'hen unloaded. As loading c·ontinues.
outcrni"ost fibers of the struCture begin to yield
some point. This yidd poinl signals the elastic li
of the structure, As, the load exceeds this limit,
structure exhibits plastic behavior, I'cflecl~d in
second (curved) portion of the eurvc, the p!as,tic
gion. The structure \\~i11 no longer return to its or
11al dimensions when the load has been releas
some residual deformation will be permanent
loading-is progressively increased, the structure
fail at'somc point (bonc' will fracturc), This poin
indic~-iled by the ultimate failure point on the CU
Plastic region
/
'" Yield
point
D
ro
S
1'-
D
1
1
1
1
/
1
c
i
Ultimate
failure
point
Energy
1
1
1
1
1
1
A
D'
Deformation
Load-deformation curve for a structure composed of
a somewhat pliable material. If a load is applied
within the elastic range of the structure (A to B on
the curve) and is then released, no permanent deformation occurs. If loading is continued past the yield
point (B) and into the structure's plastic range (B to C
on the curve) and the load is then released, permanent deformation results. The amount of permanent
deformation that occurs if the structure is loaded to
point 0 in the plastic region and then unloaded is
represented by the distance between A and D. If
loading continues within the plastic range, an ultimate failure point (C) is reached.
•
Three parameters for determining the strength of
a structure are reflected on the load-deformation
Clll-ve: 1, the load that the structure can sustain before failing; 2, the deformation ihat it cansustain
before failing; and 3, the erler¥.Y that it CaIlstore be~
fore failing. The strel1gth in terms of loaclancLdefm:"mation, or ultimate strength, is in,clicatedOllJhe
curve by the ultimate failure point. The streI,lgtl1 in
terms of energy storage is indicated by the size of
lhe area under the entire cun'e. The larger the area,
the greater the energy that build~up in tfle structure as the load is applied. The stiffness of the
structure is indicated by theslope of tl).<::.curve in
the elastic region. Thesteepei::::the slope, the stifrer
the material.
the load-deformation curve is useful for deterInining the mechanical properties of whole struc~
turessuch as a whole bone, an entire ligament or
tendqn, or a metai implant. This knowledge is helpful in the study of fracture behavior and repail~ the
response of a struetlJre to physical stress, or the er~
fect of various treatment programs. However, char~
acterizing a bone or other structure in terms o
material of \vhich it is composed, independent o
geometry, requires standardization of the tes
conditions and the size and shape of the test sp
mens. Such standardized testing is useful for c
paring the 111echanical properties of two or m
materials, such as the relative strength of bone
tendon tissue or the relative stiffness of various
terials used in prosthetic implants. More pre
units of measurement can be used when stand
ized samples are tested-that is, the load per un
area of the sample (stress) and the amount of de
mation in terms of the percentage of change in
sample's dimensions (strain). The curve generat
a stress-strain curve.
Stress)s:the load, or force, per unit area tha
velops on a plane surface within a struc:lll['e.)1
sponse'iO'exteI~ililll)' applied loads. The three ~
most commonly used for measuring stress in s
dardized samples of bone are ne\vtons per cent
ter squared (N/cm:;); newtons per meter squared
pascals (N/m 2 ,Pa); and megancwtons per m
sqmii;cd; or mega pascals (MN/m 2 , MPa).
Strain is the clef()l'mation (change in dimens
that develoP?~yithiI1.~ls:tTuctureirl"I:t::sponse to
ternallyapplied loads. The two basic types of st
are ii'near strain, which causes "a- chang~ ill
lengtl;"'6f'dlcspecimen, and shear strain, w
causes }i"'~I~~i:~¥?i~.(ll~angulari'·clationships \v
the structure. Linea-I'· strain is measured as
amoliiltofflIlear de[ormati()n (lengthening or sh
ening}"6fthesiilnple di\'ided by the sample's orig
length. It is a nondim<;nsj()llal paramel<;r expre
as a percentage (e.g:, centimeter per centime
She'~1'1:'stt'~lTI1"is measured as the all"1Q.~lIlL.of ang
ch~~!!.g,~. ,,,(Y)_i,l'-adglif _ ~~lj:~leI )'i l1gi"I"i.t}l<; pl.nne.o
terest in the sample. It is expressedil"il'actians
radian-'e(ill~llsai)proximately57.3°) (Internatio
Society of Biomechanics, 1987).
Stress and strain values can be obtained for b
by placing a standardized specimen of bone ti
in a testing jig and loading it to failure (Fig. 2
These values can then be plotted on a stress-st
curve (Fig. 2-8). The regions of this curve are s
lar to those of the load-deformation cUll/e. Load
the elastic region do not cause perll1aD.<;nL~I~J(
lion, buC6ncc:the yield point is excee.deeJ,s()ll1c
Formation L5 permanent. The strengthoftheJl1~l
inl in terms of energy' storage is repre~ent~elby
area .~II1~1.<::I·theyntire curve. The stifFness is re
sented h.ytheslope of th<? curve in the~Iasticreg
A value for stiffness is obtained by dividing
stress at ~ __point ir~_.thc clastic (straight line) porti()11
of trlC"cllI"\'e by i·he-~t~,~~·in ~t that point. This
-is
caITedtTle-nl0dlll~I~~ o!:-elastfcity (Young's modulus).
Young'S modulus (E) is d~,T;'ed from Ihe relalionship
betw';-en str"ss ('T) an~.strain(~):
value
E=<r/E
The elasticity 01" a material or the Young's modulus
E is equal to the slope of the stress (<r) and strain (E)
diagram in the clastic linear region. E represents the
stiffness of Ih" material, such Ihat Ihe higher the
elastic modulus or Young's modulus, the stiffer the
material (Ozkaya & Nordin, 1999).
Mechanical propcrties differ in the two bone
types. Cortical_bq__,!_c ~~ s~.Hfer_tlla!,! canccllQus_ bone,
withstanding greater stress but less strain before
falllll·c. C:-1-n~c~ITol-i-s -bolic-rn·-\'itl:-O--~~·~I~-~;jn lip to
50% of strains before yielding. \vhile cortical ~)(?nc
viclds and fractures when the strain exccec!s 1.5
10 2.00/0. 8cc~\i.isc---<.)r ·ils~·I)O-'·()lis---stl~ucl·ure, ca..Q~J~I.~_
IOlls~~~=_.I:..'~_s a la-j";gc c~:.pa.::itS~_ (61~ enei·gy..sto,~ge
(Ke"""ny & Hayes, 1993). The physical differen
betwecn the two bone tissucs is quantified in ter
of the apparent density of bone, which is ddined_
the mas~_.<?Ip·9ii~jI~~u£:i?,:·~~~ntin a unit of bon~ \
~~f11e (gram per cubic ccn-titnctci:Tglcc]):-Frg"Llre
depicts typical stress-strain qualities of cortical a
trabecular bone with different bone densities tes
under similar conditions. In general, it is
enough lo describe bone strength with a sin
number. A bctter way is to examine the stress-str
curve for the bone tissue under the circumstan
tested.
To better understand th" relationship of bone
other materials, schematic stress-strain curves
bone, metal, and glass illustrate the differences
mechanical behavior among these matcrials (F
2~10). The variations in stiffness are rcOecled in
different slopes of the cun'cs in the clastic regi
Metal has the steepest slope and is thLis the stiff
material.
C' .•• ---- .. -••• -------- ••••........••••........••••....
C
PlastiC region
B'
A
B"
C"
Stress-strain curve for a cortical bone sample teste
in tension (pulled). Yield point (B): point past whic
some permanent deformation of the bone sample
curred. Yield stress (8'): load per unit area sustaine
by the bone sample before plastic deformation too
place. Yield strain (8"): amount of deformation wi
stood by the sample before plastic deformation oc
curred. The strain at any point in the elastic region
of the curve is proportional to the stress at that
point. Ultimate failure point (C): the point past
which failure of the sample occurred. Ultimate stre
(C'): load per unit area sustained by the sample be
fore failure. Ultimate strain (C"): amount of defor
mation sustained by the sample before failure.
~--Standardized bone specimen in a testing machine_
The strain in the segment of bone between the two
gauge arms is measured with a strain gauge. The
stress is calculated from the total load measured.
Courtesy of Dennis R. Carter. Ph.D.
•
Apparent Densily
~
~
~
..... 0.30 glee
:::1 .
- - . 0.90 glee
CortIcal bone
- - 1.85 glee
100
iii
50
/
/
--------------
/
o.j.........,=.:...:.
.:".+"..:..:. ".:....:..:.:. :".:."
Trabecular bone
:.c
..c.;.,:..:.":..
•.c.;":..:•.:.":...".f.:.c.;".:.":..".:.."---<
15
20
25
-'<":..:.":..:..:.
. .:..:...
o
5
10
scnce of a plastic region on the stl"L'ss-strain cur
By contrasl, metal exhibits cxtensh·t: deformati
before failing, as indicalC'd b~' a long plastic reg
on the ClitYC. Bone also deforms before failing
to a rnuch lesser extent than metal. The differen
in the plastic behavior of metal and bone is the
suit or differences in microlllcchanical events
yield. Yielding in melal (tested in tension,
pulled) is caused b~' plastic rIow and the fonrwti
of plaslic slip lines; slip lines art: formed when
molecules of lhe latticc structurc of mctal dis
cate. Yielding in bone (tested in tcnsion) is caus
Strain (%)
Met
Example of stress-strain curves of cortical and trabecular bone with different apparent densities, Testing
was performed in compression. The figure depicts the
difference in mechanical behavior for the two bone
structures. Reprinted with permission from Ke,1'.'eny.
T M., & Hc1yes, \tV. C. (1993). Mechanical properri(;.ls of cortical ancl rfDoecular bone, Bone. 7, 28S·]t]4,
The clastic portion of the curve for glass und
metal is a straight linc. indicating linearly ei<:\slic behavior; virtually no yielding lakes place before the
yield point is reached. By comparison. precise testing of cortical bone hns shown thnt the elastic portion of the curve is not straight but instead slightly
curved. indicating that bone is not linearly claslic in
its bdlm'ior but yields somewhat during loading in
the elastic rcgion (Bonefield & Li, 1967)" Table 2-1
depicts the mechanical propcrtics of selectcd bio·
materials for comparison. Materials are classified as
brittle or ductile depcnding on the extent of defor·
mation before failure. Glass is a typical brittle material. and soft metal is a typical ductile material.
The cUrrerence in the amount or deformation is reOected in the fracture surfaces or the two materials
(Fig" 2-11)" When pieced togethcr aftcr fracture, lhe
ductile material will not conform to its original
shape whet'cas the brittle material will. Bone ex·
hibits more brittle or ITIOI'C ductile behavior dt>
pending on its age (younger bone being more ductile) and the rate at which it is loaded (bonc bcing
more brittle at higher loading speeds).
After the yield point is reached, glass deforms
very little before failing, as indicated by the ab·
Glass
Bone
Strain
Schematic stress-strain curves for three materials.
Metal has the steepest slope in the elastic region a
is thus the stiffest material. The elastic portion of
curve for metal is a straight line, indicating linearl
elastic behavior. The fact that metal has a long pla
tic region indicates that this typical ductile materia
deforms extensively before failure. Glass, a brittle
material, exhibits linearly elastic behavior but fails
abruptly with little deformation, as indicated by th
lack of a plastic region on the stress-strain curve.
Bone possesses both ductile and brittle qualities
demonstrated by a slight curve in the elastic region
which indicates some yielding during loading with
this region.
•
~
by dcbonding of thL' ostC'ons 4\t the cement lin
and micl'ofracturc (Fig. 2-12), while ~'idding
bone as a result of compression is indicated
I Mechanical Properties of Selected
iI Biomaterials
cracking of lhe osteons (Fig. 2-13).
Ultimate
Strength
Modulus
Elongation'
(MPa)
(GPa)
(%)
Metals
Because the structure of bone is dissimilar in
transverse and longillidinal. directions, it exhib
diffcrerl'l mechanical propenics whc.I1 loaded alo
differenl axes, a characleristic known as anisotro
Co-Cr alloy
Cast
600
220
forged
950
850
900
220
210
110
Stainless steel
Titanium
15
10
15
Polymers
Bone cement
20
2.0
2-4
,.-;
Ceramic
Alumina
300
350
<2
100-150
8-50
10-15
1-3
20-35
2.0-4,0
10-25
Biological
Cortical bone
Trabecular bone
Tendon, ligament
2-4
Adapted from Kummer, J,K, (1999) Impldn, biomillcrials In J.M.
Spivak, P.E, DiCesare, Os. Fe[dman, K,!. Ko'''.1[, A.S. RokilO, & J.D.
Zuckerrnan (Eds.). Orrhopaedic5.' A Study GlII'de (pp. 45-48). Ni:>,v
Reflected-light photomicrograph of a human corti
bone specimen tested in tension (30):). Arrows ind
cate debonding at the cement lines and pulling ou
of the osteoos. Courtesy 0; Dennis R. Carter. Ph.D.
York: lvlcGraw-Hlil.
1t - - - - - - - II
i
Ductile fracture
I
I
-.
I
Brittle fracture
~
I
,urface' of sample, of a ductile and a br;ttle
material. The broken Jines on the ductile material indicate the original length of the sample. before it
deformed. The brittle material deformed very little
before fracture.
Scanning electron photomicrograph of a human co
tical bone specimen tested in compression (30X). A
rows indicate oblique cracking of the osteons. Courtesy
Dennis R. Carter, Ph.D.
•
200
MPa
----
Strain
Anisotropic behavior of cortical bone specimens from a
human femoral shaft tested in tension (pulled) in four
directions: longitudinal (L). tilted 30" with respect to
the neutral axis of the bone, tilted 60", and transver
(T). Data from Frankel, V.H., & Burstein, A.H. (1970).
lhopaedic Biomech<1nlcs. Philadelphia: Lea & Febiger.
•
f·7jgure 2-14 shows the variations in strength and
stiffness for cortical bone samples from a human
remoral shah, tested in tension in four directions
(Frankel & Burstein c 1970; Carterc 1978). The values
for both parameters are highest for the samples
loaded in the longitudinal direction. Figures 2·9 and
2·15 show trabecular bone slI-cngth and stiffness
tesled in two directions: compression and tension.
Trabecular or cancellous bone is approximately 25%
as dense, 5 to look; as stiff, and five times as ductile
as conical bone.
Although lhe relationship bel ween loading patlerns and the mechanical properties of bone
throughout the skeleton is extremely complex, it
generally can be said that bone strength and stiffness are greatest in the dircction in which daily
loads arc most commonly imposed.
10
B
6
4
2
0-1----1----1----1---+--"'"-1
o
BiOlnechanical Behavior
of Bone
The mechanical behavior or bone-ils behavior
under the innuence of forces and moments-is affected by its mechanical properties, its geometric
characteristics, the loading mode applied, direclion of loading, rale of loading, and frequency of
loading.
2
4
6
B
Tensile strain (%)
Example of tensile stress-strain behavior of trabec
lar bone tested in the longitudinal axial direction
the bone. Adapted from Gibson, L.1., & As/lby. M.F.
(1988). Cellular Solids: Structure and Properties. New
York: Pergamon, Press.
pendicular to the applied load (Fig. 2-17). Under
stnl.c:ture lengthens and narrows
Clinically, fractures produced by tensilelqad
are usually seen in Dones with a large proportio
cancellous bone. Examples are fractures of the b
or thCfifth metatarsal adjacent to the attachm
of the pe-roneus brevis tendon and fractt,lres
the calcaneus adjacent to the attachment of
AchUlqs tendon. Figure 2-18 shows a tensile frac
through the calcaneus; intense contraction of
triceps surae muscle produces abnormall,Y high
sile loads on the bone.
sil~l(),l(ling, the
..
Compression
Tension
•
•
Shear
I'~.
Compression
I
I
i
eli
C..J
Torsion
~L.-.-
1
Bending
Combined
loading
_
Schematic representation of various loading modes.
DUling compressive loading, equal ~nd opposite lO
are applied toward the surface of the structure
compressive stress and strain. result inside the st
ture. Compressive stress can be thought of as m
small forces directed into the surface of the stluct
Maximal compressive stress occurs on a plane
pendicular to the applied load (Fig. 2-19). Under c
pressive loading, the structure shortens and widen
Clinically, COin pression fractures are commo
found in the vertebrae. which are subjected to h
compressive loads. These fractures are most o
seen in the elderly with osteoporotic bone tis
Figure 2-20 shows the shortening and widen
BONE BEHAVIOR UNDER VARIOUS
LOADING MODES
Forces and moments can be applied to a structure
in various directions, producing tension, compression, bending, shear, torsion, and cornbinecl loading
(Fig. 2-16). Bone in vivo is subjected to all of these
loading modes. The folknving descriptions of these
modes apply to structures in equilibrium (at rest or
moving at a constant speed); loading produces an
internal, deforming efFect on the structure.
Tension
During tensile loading, equal and opposite loads are
applied olltward from the surface of the structure,
ancCtensile stress and strai"n result inside the structure. Ten'sile stress can be thought of as manv small
forces directed· <.nva:v from the -;'urface of th~ stlJIClure. Maximal tensile stress occurs on a plane per-
Tensile loading.
Tensile fracture through the calcaneus produced by
strong contraction of the triceps surae muscle during
a tennis match. Courtesy of Robert A. Winquisr, lA.D
that takes place in a human vertebra subjected to a
high compressive load. In a joint, compressive loading to failure can be produced by abnormally
strong contraction of the surrounding muscles. An
example of this effect is presented in Figure 2-2 t;
bilateral subcapilal fractures of the Femoral neck
Compression fracture of a human first lumbar ver
bra. The vertebra has shortened and widened.
•
were sustained by a patient undergoing electroc
vulsive therapy; strong contractions of the mus
around the hip joint compressed the femoral h
against the acetabulum.
..
Compressive loading.
Shear
During shear loading, a load is applied parallel to
surface of the structure, and shear,stress and st
result inside the structure. Shear stress can
thought of as many' small forces acting on the
face of the structure on a plane parallel to the
plied load (Fig. 2-22). J\ structure subjected t
shear load deforms internaH.v in an angular-rllan
right angles on a plane surface within the stnlc
Force
L
Before loading
;~
1Dfm
I
I
_
Bilateral subcapital compression fractures of the
femoral neck in a patient who underwent electrocon·
vulsive therapy.
e
f
become obtuse or acute (Fig. 2-23'). \'Vhel~e\'er a
structure is subjected to tensile or compressivelqading, sflear stl:ess is produced. -Figli;·c 2-24 iIllls11:a'tcs
angtiliii· deformatiorl in sfi~lctllrcs subjected (0 these
loading modes. Clinically, sh<;ar fractures are most
ohen seen in cancel lOlls bone.
Human «clllit 'corticai bone exhibits different values for ultimate stress uncleI' compressive, tensile,
and shear loading. Cortical bone can withstand
greater stress in compression (approximately 190
Mpa) than in tension (approximately 130 !\Ilpa) and
greater stress in tension than in shear (70 r\'lpa). The
Under shear loading
When a structure is loaded in shear, lines original
at right angles on a plane surface within the struc
ture change their orientation, and the angle beco
obtuse or acute. This angular deformation indicat
shear strain.
.-----------------
daslicily (Young's modulus) is approximately
GPa in longitudinal or axial loading and appro
1l1alcly I I GPa in transverse loading. Human
becular bone values for testing in compression
approximately' 50 ivlpa and arc reduced lO appr
***
D
o
<> 0
1
Unloaded
I
F.
~
'---------
_
Shear loading.
1
-~
Under
lensile
loading
CJ
<>
***
Under
compressive
loading
The presence of shear strain in a structure loaded
tension and in compres.sion is indicated by angula
deformation.
Cross-section of a bone subjected to bending, showing distribution of stresses around the neutral axis.
Tensile stresses act on the superior side, and compressive stresses act on the inferior side. The stresses are
highest at the periphery of the bone and lowest near
the neutral axis. The tensile and compressive stresses
are unequal because the bone is asymmetrical.
lateral roentgenogram of a "boot top" fracture
duced by three-point bending. Courtesy of Robert
\J1/inquist, M, D.
mately 8 Mpa iF loaded in tension. The modulus or
elasticity is low (0.0-0.4 GPa) and dependent on the
apparent density of the trabecular bone and direc~
lion of loading. The clinical biomechanical consequence is that the direction of compression failure
results in general in a stable fracture, while a fracture initiated by..' tension or shear ma!' have catastrophic consequences,
....
1_/
..
A_-------.JO
....
..
B_-----.JO
1_/_ _
Two types of bending. A, Three-point bending.
B, Four-point bending.
Bending
In bending, loads are applied to a structure
manner that causes it to bend ~\bout an,axis. \V
a bone 'is loaded in bending, it is subjected
combination of tension and compression. Te
stresses and strains act on one side of the ne
axis!._'111d compressive stresses and strain?,Ic;t on
oth~r side (Fig. 2-25); there are I}ostresse,:str~liI!?along tt;enelltral axis. The ~lagnitllde o
stresses is proportional to their distance from
neutral axis of the bone. The farther the stresse
from the neutral axis, the higher their magnit
Because a bone structure is asymmetrical,
stresses may not be equally distributed.
Bending may be produced by thl'qe forces (th
point bendiilg) Ol:.Xour forces (roUI'~point bend
(Fig. 2-26). Fractures produced b)' both type
bending are commonly observed clinically, par
lady in the long bones.
Three-point bendirlgtakesplace when t
forces acting on a structure pn)c!uce two equal
mcnts, each being the product of one of the t\\'
ripfle!~;;liforces and its perpcndicular distance.
the axis of rotation (the point at which the mi
Forceis applied) (Fig. 2-26;\). IF loading conti
to the )-'iele! point, the structure, if homogene
symmetrical, and with no structural or tissue
,"''''
~
Fatigued muscle
o
,
~
~~
_"~ ~
"
~
~ ~ ~"_c
feet, will break at the point of application of the
middle force.
A typict;~Jthree~pointl)eIl(ling fracture is tht: "boot
top"Jracturesustainedb:vskiers. In the "boot top"
fracture shown in Figure 2-27, one bending moment
acted on the proximal tibia as the skier fell forward
over the top of the ski boot. An equal moment, produced by the fixed foot and ski, acted on thc distal
tibia. As the proximal tibia was bent fonvard, tensile
stresses and strains acted on the posterior side of
the bone and compressive stresses and strains acted
on the anterior side. The tibia and fibula fractured
at the top of the boot. Because adult bone is weaker
in tension than in compression, failure begins on
the side subjected to tension. Because immature
hone is n10r~ ductile, it nUl)' fail first in compression, and a buckle fracture may result on the compressive side (Flowchart 2~ I).
Four~point bending takes place when two fo
couples acting on a stl'uclurc produce two eq
moments. A force couple is formed when two par
lei forces or equal magnitude but opposite direct
are applied to a structure (Fig. 2-28;\). Because
magnitude of the bending moment is the sa
throughout the area between the two force coupl
the structure breaks at its weakest point. An exa
ple of a FourNpoint bending FraclUre is shown in F
ure 2-28B. A stifT knee joint was manipulated inc
rectl y' during rehabilitation of a palient with
postsurgical infected femoral fracture, During
A
•
A, During manipulation of a stiff knee joint during
fracture rehabilitation, four-point bending caused
the femur to refracture at its weakest point, the
original fracture site. B, Lateral radiograph of the
fractured femur. Courtesy of Kaj Lundborg, M.D.
Cross-section of a cylinder loaded in torsion, showing
the distribution of shear stresses around the neutral
axis. The magnitude of the stresses is highest at the
periphery of the cylinder and lowest near the neutral
axis.
manipulation. the posteJ"ior knee joint capsule and
tibia Formed one force couple and the fClTlOral head
and hip joint capsule formed the other. As a bending
moment was applied to the femur, the bone failed at
its \vcakest point, the original fracture site.
Torsion
In torsion, a load is applie~l to a strl1c~.lII·c _in a mannel' that causes i"l to l\\;j'st about an axis, and a
lO~'gue--(or ornament) is produced within the -struclure.--\Vhen a structure is loaded in torsion, shear
61
, .
, Ii'
<J)
,,, ,!'
~~
@.-
Shear
,:
Compression
,i<?
.......
1
'-.-'
stre:;ses arc distributed over the entire structure
in b~n'ding, the magnitude of these stresses is
ponional to their distance from the neutral
(Fig. 2·29). The farther the stresses arc from
neutral axis, the higher their magnitude.
Under torsional loading, maximal shear stre
act on plan~sp~~-;'alleland~perpen(ficular·to-then
tl"al axis of the structure. 'In addition, max"fillal
sile aild comp'ressi\'c stresses act on a plane dia
nal to the neutral a.'\is of the structure. Figure
illustrates these planes in a small segment of b
loaded in torsion.
The fracture pattern for bone loaded in tor
suggests thal the bone fails first in shear, with
formation of an initial crack parallel to the neu
axis of the bone. A second crack usually fo
along the plane or maximal tensile stress. Suc
pattern can be seen in the experimentally produ
torsional fracture of n canine femur shown in
ure 2-31.
Tension
I
,;
Experimentally produced torsional fracture of a c
nine femur. The short crack (arrow) parallel to th
neutral axis represents shear failure; the fracture
at a 30'" angle to the neutral axis represents the
plane of maximal tensile stress.
Schematic representation of a small segment of bone
loaded in torsion. Maximal shear stresses act on
planes parallel and perpendicular to the neutral axis.
Maximal tensile and compressive stresses act on
planes diagonal to this axis.
Combined loading,
Allhough each loading mode has been conside
separately, living bone is seldom loaded in one m
only. Loading of bone in vivo is complex for
principal reasons: bones .~.re ~onstantly subjecte
multiple indeterminate loads and their geome
structure is irregulm: In vivo mcasurement of
strains on the antcrol11cdial surface of a hum
adull tibia during walking and jogging dcm
SlrateS the comph.:.xil.V 01" the loading patterns during these cOl11mon ph.'·siological activities (Lanyon
el aI., 1975). Stress values calculated from these
strain measurel~lents by Carter (1978) showed thal
during normal walking, the strc.i.~es wcre COlllP.I"CSs.ivc dllring heel strike, tensile during- thc--s_t-~!.nce
ph~y~·,·,indi1¥aill--c()il·~prcssive-.~Iu'ri,~,lg- pLlsl.1~off (Fig.
2-3"2A). VaI"ucs for shear stress \\'cn.:.~ relatively high
in the later portion of the gait c)ldc, denoting significant torsional loading. This torsional loading
\vas associated with external rotation of the tibia
during stance ancl push-ofr.
Dul"ing jogging. the stress pattern was quite difrerent (Fi-g. 2:328). The COlJlP.~·.~~.~~c_slrcssPFcdominatin!! at lac strike was followed bv high tensile
stress~di.iring r>ll-sh-ofr. The sllenr slrcss~:vi~~--'o\V
tfil'oughout 'tlle strich:\ denoting 111ini;lwl torsional
loadhig"pl~o(lllced by slight e~'\tern~" and inlcrm~l 1'0tatiOll or-the tibia in an-ahernating pallcrn. The increase--fll specd from slow walking LO jogging in-
crcas~d
both the str~ss and the strain Oil the lib
(Lan.'·on ct aI., 1975). This increase in strain w
grt,,:atcr speed was confirmed in studies or locom
lion in sheep, which ckmonslratcd a fivefold
crcase in slrain values from siD\\" walking to f
lrotting (Lanyon & Bourn, 1979).
INFLUENCE OF MUSCLE ACTIVITY ON STRESS
DISTRIBUTION IN BONE
When bone is loaded in vivo, the contraction of t
muscles attached to the bone alters dlC stress dist
bution in [he bone. This muscle contraction d
cr<.:ases 01· eliminates [ensile stress on the bone
producing compressi\'c SlI-CSS thal neutralizes it
ther partially or totall~·.
The effect of muscle contraction can be ill
trated in a tibia subjected to three-point bendin
Figure 2~33A represents Lhe leg of a skier who
falling: forward, subjecting the tibia Lo a bendi
Jogging (2.2 m/sec)
12
10
Walking (1.4 m/sec)
4
Stress
Tensile
Compressive
Shear (external rotation)
3
2
~
z
1
;;; 0 +''<,--+-!----o:..-.--+,__+
~"'
-',
{jj
!
\ .....;,
4·
A
;i ••••••••
\.
3
rn
a.
~
I
1
i
~
~
6
4
~
iii
__.-;~'c...,...+-
'
..j
Tensile
Compressive
Shear (external rotation)
Shear (internal rotation)
8
2
...i
-~'..
.:;
... i
"'
\ .. i "
r
h.
l,\
HS
FF
HO
HO
TO
2·
(1975). Bone deformation recorded in vivo (rom srrain
gauges c1tr<elched to the human tibial sharr. ACla Orlhop
Scand, 46. 256. Figure coortesy of Dennis R. Carler, Ph.D
"'
i
i
"'\ ii
\i
s
A. Calculated stresses on the anterolateral cortex of a
human tibia during walking. HS, heel strike; FF, foot
flat; HO, heel-off; TO. toe off; S, swing. Calwlated from
Lanyon, L.E.. Hampson. W'.G.J., Goodship. A.E.. et af.
................
i
O' ,
\·······1··../ /
1\
B
4 .l---=-E--=~'-:':::-----TS TS·TO
8, Calculated stresses on the anterolateral cortex of a
human tibia during jogging. T5, toe strike; TO, toe o
Calwfared [rom Lanyon. L.E.. Hampson. W.G.)., Goodship
AE., et al. (1975). Bone deformation recorded in vivo fro
strain gaoges aClacIJed ro the human tibial Shaft. Acta
Orthop 5cand. 46. 256. Figure courtesy 01 Dennis R. Carte
Pll.D.
, ,-, ,
,,
•
""
r:~
, ,
,, ''
:
I
B
A. Distribution of compressive and tensile stresses in
a tibia subjected to three-point bending. B, Contraction of the triceps surae muscle produces high compressive stress on the posterior aspect, neutralizing
the high tensile stress.
moment. High tensile stress is produced on the postcrior aspect of the tibia, and high compressive
stress acts on the anterior aspect. Contraction of the
triceps surae muscle produces great compressive
stress on the posterior aspect (Fig. 2-338), neutralizing the great tensile stress and thereby' protecting
the tibia from failure in tension. This muscle contraction may result in higher compressive stress on
the anterior surface of the tibia and thus protect the
bone from failure. Adult bone can usuall~/ withstand
this stress, but immature bone, which is weaker,
ma.y fail in compression.
Muscle contraction produces a similar effect in
the hip joint (Fig. 2-34). During locomotion, bending moments are applied to the femoral neck and
tensile stress is produced on the superior cortex.
Contraction of the gluteus medius muscle produces compressive stress that neutralizes this tensile stress, with the net result that neither compressive nor tensile stress acts on the superior
cortex. Thus, the muscle contraction allows the
femoral neck to sustain higher loads than would
otherwise be possible.
STRAIN RATE DEPENDENCY IN BONE
Because bone is a viscoelastic material, its biomechanical behavior varies with the rate at which the
bone is loaded (Le., the rate at which the load is
plied and removed). Bone is stifTer and sustain
higher load to failure when loads are applied
higher rates. Bone also stores more energy be
failure at higher loading rates, provided that th
rates are within the physiological range.
The in vivo claily' strain can vary considerab
The calculated strain rate for slow walking is 0.
per second, \vhile slow running displays a st
rate of 0.03 per second.
In general, when activities become more stre
ous, the strain rate increases (Keaveny & Ha
1993). Figure 2-35 shows cortical bone behavio
tensile testing at different physiological strain ra
As can be seen from the figure, the same chang
strain rate produces a larger change in ultim
stress (strength) than in elasticit.\' (Young's mo
lus). The data indicates that the bone is appr
mately 30(/0 stronger for brisk walking than for s
c"~,~
,/
:,
,,:
,
,
,
,
E
,
,,
,
;,
,
'
'
'
'
:
"
Stress distribution in a femoral neck subjected to
bending. When the gluteus medius muscle is relax
(top), tensile stress acts on the superior cortex an
compressive stress acts on the inferior cortex. Con
traction of this muscle (bottom) neutralizes the te
sile stress.
400
1500/sec
300
300/sec
"
I~
l::l
~--- O.lIsec
200
'I"
- - - - - O.Ollsec
iii
Brisk walking
O.OOl/sec
100
Slow walking
o'l'---+---+---+------i
0.0
0.5
1.5
2.0
Rate dependency of cortical bone is demonstrated at
five strain rates. 80th stiffness (modulus) and
strength increase considerably at increased strain
rates. Adapted from McElhaney, J.H. (1966). Dynamic
response of bone and muscle tissue. J Appl Physio1, 2},
I,
/23/-/236.
III
walking. At very high strain rates (> I per second)
representing impact trauma, the bone becomes
more brittle. In a full range of cxperimentaltesting
for ultimate tensile strength and elasticity of corti, cal bone, the strength increases by a factor of three
and the modulus by a faclor of two (Keaveny &
Hayes, 1993).
The loading rale is clinically significant because it
innuences both the fracture paUern and the amount
of soft tissue damage at fracture. \Vhen a bone fractures. the stored energy is released. At a low loading
rate, the energy can dissipate through the formation
of a single crack; the bone and soft tissues remain
relatively intact, with little or no displacement of the
,. bone fragments. At a high loading rate, however, the
;'·',greater energy stored cannOt dissipate rapidly
:~ enough through a single crack, and comminution of
bone and extensive soft tissue damage resull. Figure
. 2-36 shows a human tibia tested in vitro in torsion
at a high loading 1'4ue; numerou:j bone fragments
;;J'
were produced, and displacement of the fragme
was pronounced.
Clinically, bone fractures fall into three gene
categories based on the amount of energy releas
at fracture: low-energy!, high-energy, and VCI)' hi
energy. A low-energy fracture is exemplifled by
simple torsional ski fracture; a high-energy fract
is often sustained during automobile accidents; a
a very high-energy fracture is produced by v
high-muzzle velocity gunshot.
FATIGUE OF BONE UNDER
REPETITIVE LOADING
Bone fractures can be produced b:v a single lo
that exceeds the ultimate strength of the bone
by repeated applications of a load of lower mag
tude. A fracture caused by a repeated load appli
tion is called a fatigue fracture and is lypically p
duced either by few repetitions of a high load or
many repetitions of a relatively normal load (C
Study 2-1).
The interplay of load and repetition for any m
te!"inl can be plotted on a fatigue curve (Fig. 2-3
For some materials (some metals. for exampl
thc fatiguc curve is asymptotic, indicating tha
the load is kept below a certain level, theoretica
the material will remain intact no matter h
many repetitions. For bone tested in vitro,
curve is not asymptotic. \,Vhen bone is subjected
repetitive low loads, it may sustain microfractur
Testing of bone in vitro also reveals that bone
tigues rapidly when the load or deformation
proaches its yield strength; that is, the number
repetitions needed to produce a Fracture dim
ishes rapidly.
In repetitive loading of living bone, the fatig
process is affected not only by the amount of lo
and the number of repetitions but also bv
number of applications of the load within a giv
time (frequency of loading). Because living bon
self-repairing. a fatigue fracture results only wh
the remodeling process is outpaced by the fatig
process-that is, when loading is so Frequent t
it precludes the remodeling necessary to prev
failure.
Fatigue fractures are llsually sustained dur
continuous strenuous physical activity, wh
causes the muscles to become fatigued and redu
.' their ability to contracl. As a result, the.v are l
able to store energy aI~d thus to neutralize
stresses imposed on the bone, The resulling al
ation of the stress distribution in the bone cau
Human tibia experimentally tested to failure in torsion at a high loading rate. Displacement of the numerous fragments was pronounced.
B"ne Overloading
vent failure. iVluscie fatigue occurred as a result of the abno
mal loading pattern and the intensive uaining, It affected th
23-year-old military recruit was exposed to an intensive
" , ' heavy physical training regime that included repetitive
A
continuous crawling in an awkward position for several
weeks (Case Study Fig. 2-1-1 A). The repeated application of
loads (high repetitions) and the number of applications of a
load during a short period of time (high frequency of loading)
muscle function in the neutralization of the stress imposed,
leading to abnormal loading and altered stress distribution
(Case Study Fig. 2-1-1 B).
After 4 \-veeks of strenuous physical activity, the damage
accumulation from fatigue at the femoral shaft lead to an
oblique fracture.
surpassed the time for the bone remodeling process to pre-
Case Study Figure 2-1-1A. Abnormal loads at the femoral shaft occurred.
,'.
".3
m
Injury
Repetition
The interplay of load and repetition is represented
on a fatigue curve.
•
abnormally high loads to be imposed, and n fatigue
damage accumulation occurs that Illav lead to
a fracture. Bone may fail on the tensil~ side, on
the compressive side, or on both sides. Failure on
the tensile side resulls in a transverse crack, and the
bone proceeds rapidly to complete fracture. Fatigue
fnlctures on the compressive side appear to be pro-
II ::1
'E
0006
I~
0 004
l~
Ii g~
·
I
tI)
0.002
o Compression
• TenSion
Miles
10
100
1000
I
•
I
!
•
vlgorous,-:":::!::::""~~S~'b-'<:::::::-;:;-~
exercise
Running -----------===--:::-
Walking
0.0001+0-0--1---+---+----+---!------I
1~
1~
1~
Number 01 Cycles
"i
lam
I
--------
.
I
I
I
'
,i
i
i
Fatigue testing showing the number of cycles (x-axis)
and strain range (y-axis) expressed as stress rangel
modulus in human cortical bone specimens loaded in
tension and compression. Typical strain ranges are
shown for walking, running. and vigorous exercises.
Note that resistance to fatigue fracture is greater in
compressive loading. Ten miles represent approxj·
mately 5,000 cycles, corresponding to the number of
steps running during that distance. Aclapared from
Carrer. D.R., Cater. W.E., Spengler, O.M.. Frankel, \l.H.
(J 98 1). Fatigue behavior of adtilt cortical bone: the influence of mean srrc1ilJ anel strain range. Acra Orthop Seane!.
52.48/-490.
J.-----------------
duced more slowl~'; the remodeling is Icss casily o
paced by the fatiguc process and the bone may
proceed to complete fracture.
This theol)1 of muscle fatigue as a cau~e of
tigue fracture in the lower extremities is outlin
in the schema in Flowchart 2-1 on p. 41.
Figure 2-38 shows typical strain ranges for
man femoral cortical bone during different act
tics and distances. Resistance to fatiguc behavio
great.er in compression than in tension (Keaveny
Hayes, 1993). On average, approximately 5,000
clcs of experimental loading correspond to
number of steps in to miles of running. One m
lion cycles corresponds to approximately 1,
miles. A total distance of less than 1,000 m
could cause a fracture of the cortical bone tiss
This is consistent with stress fractures repor
among military recruits undergoing strenu
training of up to 1,000 miles of nll1ning ove
short period of timc (6 weeks). Fracturcs of in
vidual trabeculae in cancellous bone have been
served in postmortem hUlllan specimens and I
be caused by fatigue accumulation. Common s
arc the lumbar vcnebrae, the femoral head, a
the proximal tibia. It has been suggested that th
fractures may playa role in bone remodeling
well as in age-related fractures, collapse of s
chondral bone, degeneraLive joint diseases, a
other bone disorders.
INFLUENCE OF BONE GEOMETRY ON
BIOMECHANICAL BEHAVIOR
The geometrv of a bone greatl\' influences its m
chanical bel~avior. In Le~lsion' and compressi
the load to failure and the stiffness arc prop
tional to the cross-sectional area of the bone. T
larger the area, the stronger and stiffer the bo
In bending, both the cross-sectional arca and
distribution of bone tissue around a neutral a
affcct the bone's mechanical behavior. The qu
tity that takes inLO account these two factors
bending is called the area moment of inertia
larger moment of inertia results in a stron
and stiffer bone. Figure 2-39 shows the in
ence of the arc,) moment of inenia on the l
to failure and the stiffness of three rectangu
Slruclures thal have the same area but differ
shapes. In bending, beam III is thc stillest of
lhree and call withstand the highest load beca
the greatest amount of material is distribuLed a
distance from the neulral axis. For rectangu
cross-sections, the formula for the area momcn
- - -
- ----,1--
4 ){ 1
2x2
1x 4
I
II
III
Three beams of equal area but different shapes subjected to bending. The
area moment of inertia for beam I is 4/12; for beam II, 16112; and for beam
111,64/12. Adapted hom Franke', VH .. & Burstein, AH. (970). Orthopaedic Biomechanics. Philadelphia: Lea & Febiger.
inertia is the width (8) multiplied by the cube of
the height (1-1') divided by 12:
B· H'
12
Because of its large area moment of inertia. bean"'!
III can withstand four times more load in bending
limn can beam I.
A third factor, the length of the bone, influences
the strength and stillness in bending. The longer the
bone, the greater the magnitude of the bending moment caused by the application of a force. ,In a rectangular structure, the magniwde of the stresses
produced al the point of application of Ihe bending
moment is proponional to lhe length of the StI1.IClure. Figure 2-40 depicts the forces acting on two
beams with the same width and height but different
lengths: beam B is twice as long as beam A. The
bending moment for the longer beatn is twice that
for the shorter beam; consequently, the stress mag~
nitudc throughout the beam is twice as high. Because of their length, the long bones of the skeleton
are subjected to high bending moments and. therefore, to high tensile and c01npressive stresses. Their
tubular shape gives them the ability to resist bending moments in all directions. These bones have a
large area moment of inertia because much of the
bone tissue is distributed at a distance from the neutral axis.
l
The factors that affect bone strength and stiff
in torsion are the same that operate in bending:
cross-sectional area and the distribution of bone
sue around a ncutral axis. The quantity that ta
into account these two factol's In torsional load
is the polar moment of inertia. The larger the p
moment of inertia. the stronger LInd stiffer the b
Figure 2-41 shows distal LInd pl'oximnl cr
sections of a tibia subjected to torsional loading
though the proximal section has a slightly sma
bony area thun docs the distal section, it has a m
higher polar moment of inertia because much o
bone tissue is distributed at a distance from the
tral axis. The distal section, while it has a la
bony area. is subjected to much higher shear st
bccause much of the bone tissue is distributed c
to the neutral axis. The magnitude of the sh
stress in the distal section is approximately do
that in the proximal section. Clinically, torsi
fractures of the tibia commonly occur distnlly.
When bone begills to heal after fracture, b
vessels and connective tissue from the periost
migrate into the region of the fracture. formin
cuff of dense fibrous lissue, or callus (woven bo
around the fracture site. stabilizing that area (
2-42A). The callus significantly increases the
and polar moments of inertia. thereby increa
the strength and stiffness of the bone in ben
and torsion during the healing period. As the r
Stress
~-"'~~-<:;magnilude
-r-_I-L---...
-L
•
S
I
Stress
magnitude =, 25
2L _____
- - - - - - 2L
----.I
as
as
bending moment. Hence, the stress magnitude
throughout beam B is twice as high. Adapted from
VH., & Burstein, A.H. (7970), Orthopaedic Bio·
mechanics. Philadelphia: Lea & Febiger.
A, Early callus formation in a femoral fracture fixe
with an intramedullary nail. B, Nine months after
jury, the fracture has healed and most of the callu
cuff has been resorbed. Courtesy of Robert A. vVinqu
MD
,,
,,,
,
,,,
,,
,,
,,
,
,,~l.,;_
Distribution of shear stress in two cross-sections of a
tibia subjected to torsional loading. The proximal
section (A) has a higher moment of inertia than does
the distal section (B) because more bony material is
distributed away from the neutral axis. Adap(ed from
Franke!, VH., & Burstein, AH, (1970;' Orthopaedic Bio·
mechanics. Philadelphia: Lea g Febiger.
ture heals and the bone gradually regains its norm
strength, the callus cuff is progressively resor
and the bone returns to as ncar its normal size a
shape as possible (Fig. 2-4213).
Certain surgical procedures produce defects t
greatly weaken the bone, particularly in torsi
These defects fall into two categories: those wh
length is less than the diameter of the bone (str
raisers) and those whose length exceeds the bone
ameter (open section defects).
A stress raiser is produced surgically' when a sm
piece of bone is removed or a screw is inserted. B
strength is reduced because the stresses impo
during loading are prevented From being distribu
evenly throughout the bone and instead become c
centrated around the defect. This defect is analog
to a rock in a stream, which diverts the watel~ p
ducing high water turbulence around it. The we
ening effect of a stress rqiser is particularly mar
under torsional loading; the total decrease in b
strength in this loading mode can reach 60~o.
c-::-<........) Empty screw hole
_ -+-... Screw in place
125
_:.~rew removed ~l_te~..-
100
~
,2
/ _.::....---
75
[;;
~
0>
<;;
c
w
50
~-.
.,..---'y
.
/'
25 <I"
raiser effect produced by the screws and by
hol('s without screws had disappeared comple
because the bone had ren1odclcd: bone had b
laid down around the screws to stabilize them,
the empty screw holes had been filled in with b
In femora from which the screws had heen
moved immediately before testing, however,
energy storage capacity of the bone decreased
50 0hi, mainly because the bone tissue around
screw sustained microdamage during sere\\'
moval (Fig, 2-43),
a
An open seclion defect is a discontinuity in
345
2
6
7
B
bone caused by the surgical removal of a piec
Weeks
Effect of
~crews
and of empty screw holes on the en-
ergy storage capacity of rabbit femora. The energy
storage for experimental animals is expressed as a
percentage of the total energy storage capacity for
control animals. When screws were removed immedi-
ately before testing, the energy storage capacity de
creased by 50%. Adapted from Bur51!.>in. A.H., et al.
4
(1972), Bone strength: The effect of sere"v floles. J Bone
Joint Surg. 54A, ] I i13.
bone longer than the bone's diameter (e.g.• by
clilling of a slot during a bone biopsy). Because
ollter surface of the bonc's cross-section is no lo
continuous. its abilit!1 to resist loads is allered,
ticularly in torsion.
In a normal bone subjected to tOl'sion, the s
stress is distributed throughout the bone and ac
resist the torque. This stress pattcl-n is illustrate
the cross-section of a long bone shown in Fi
2-44A. (A cross-section with a continuous Ollter
face is called a closed section.) In a bone wit
Contro
Burstein and associates (1972) showed the effect
of stress rabel's produced by' screws and by empty
screw holes on the energy storage capacity of rabbit bones tested in torsion at a high loading rate.
The irnmediatc effect or drilling a hole and inserting a screw in a rabbit femur was a 74% decrease
in energy stOl'age capacily. After 8 weeks, the stress
Open section
Deformation
Stress pattern in an open and closed section under
torsional loading. A, In the closed section, all the
shear stress resists the applied torque. B, In the open
section, only the shear stress at the periphery of the
bone resists the applied torque.
Load-deformation curves for human adult tibiae
tested in vitro under torsional loading. The cont
curve represents a tibia with no defect; the open
tion curve represents a tibia with an open sectio
fect. Adapted from Fr(lf)k~/. v.H., & Burstein, A.H. (1
Orthopaedic BiomechaniCS. Philadelphia: Leel & Febig
graft was removed for use in an arthrodesis of the hi
A fe\v weeks after operation, the patient tripped whi
twisting and the bone fractured through the defect.
Bone Remodeling
Bone has the ability to remodel, b)' altering its siz
shape, and structure, to meet the mechanical d
mands placed on it (Buckwalter et aI., 1995). Th
phenomenon, in which bone gains or loses cance
lous and/or cortical bone in response to the level
stress sustained, is summarized as \'VqJJr's la
which states that the, remodeling of bone is infl
enced and modulated by mechanical stress
(Wolff, 1892):
.
Load on the~,keleton can b~,,,~lc.l.:(?I""l~plished by e
tht:;,I.:,IX1~~_?(.:lc,activily or gravity. A positive correlatio
exists b~twcen bOI1e mass and body we-ight. A gr~at
body weight has been associated \vitha larger bo
mass (Exner et al., 1979). Conversely, a prolong
condition ohveightlcssness, such as that expedenc
during space travel, has been found to result in d
creased bone mass in weight-bearing bones. Astr
nauts experience a fast loss of calcium and co
sequent bone loss (Rambaut & Johnston, 197
\Vhedon, 1984). These changes are not complete
reversible.
~'------I
I
I
Normal
A patient sustained a tibial fracture through a surgically produced open section defect when she tripped
a few weeks after the biopsy.
..----------------open section defect, only the shear stress at the periphery of the bone resists the applied torque. As the
shear stress encounters the discontinuity, it is
forced to change direction (Fig. 2-44B), Throughout
the interior of the bone, the stress nms parallel to
the applied torque, and the amount of bone tissue
resisting the load is greatly decreased.
In torsion tests in vitro of human adult tibiae, an
open section defect reduced the load to failure and
energy storage to failure h,v as much as 90'10. The deformation to failure was diminished by! approximately 70% (Frankel & Burstein, 1970) (Fig. 2-45).
Clinically, the surgical removal of a piece of bone
can greatly \veaken the bone, particularly in torsion.
Figure 2A6 is a radiograph of a tibia from which a
Deformation
Load-deformation curves for vertebral segments L5
to L7 from normal and immobilized Rhesus monkeys
Note the extensive loss of strength and stiffness in
the immobilized specimens. Adapted from Kazarian,
L.L.. g Von Gierke, H.E. (1969) Bone loss as a result of
immobilization and chelation, Preliminary results in
ivlacaca mulatta. (lin Orthop. 65. 67.
Bone Remodeling
30-year-old man who had a surgical removal of an
A
..
ulna plate after stabilization of a displaced ulnar
fracture, Figure 2-48 shows anteroposterior (A) and lateral
(8) roentgenograms of the ulna after late plate removal.
The implant is used to stabilize the fracture for rapid
healing. However, in situations such as this, the late plate
removal decreased the amount of mechanical stresses
necessary for bone remodeling. It is of concern when the
plate carries most or all of the mechanical load and re·
mains after fracture healing. Thus, according to Wolff's
law, it will promote localized osseous resorption as a result of decreased mechanical stress and stimulus of the
bone under the plate, resulting in a decrease in strength
and stiffness of the bone.
Disuse or inactivity has deleterious effects on
the skeleton. Bcd rest induces a bone mass decrease of approximately' I (Ve- per week (Jenkins &
Cochran, 1969; Krolner & Toft, 1983). In partial or
total immobilization, bone is not subjected to the
usual mechanical stresses, which leads to resorp-
Roentgenogram of a fractured femoral neck to which
a nail plate was applied. loads are transmitted from
the plate to the bone via the screws. Bone has been
laid down around the screws to bear these loads.
Anteroposterior (A) and lateral (B) roentgenograms
of an ulna after plate removal show a decreased
bone diameter caused by resorption of the bone under the plate. Cancellization of the cortex and the
presence of screw holes also weaken the bone. Courtesy of Marc Marrens, 1\.1. D.
tion of the periosteal and subperiosteal bone and a
decrease in the mechanical properties of bone
(Le., strength and stiffness). This decrease in bone
strength and stiffness was sho\vn b:v Kazarian and
Von Gierke (1969), who immobilized Rhesus monkeys in full-body casts for 60 days. Subsequent
compressive testing in vitro of the vertebrae from
the immobilized monke.\!s and from controls
showed up to a threefold decrease in load to failure
and energy storage capacity in the vertebrae that
had been immobilized; stiffness was also significantly decreased (Fig. 2-47).
An implant that remains firmly" attached to a
bone after a fracture has healed may also diminish
the strength and stiFfness o( the bone. In the case
of a plate fixed to the bone with screws, the plate
and the bone share the load in proportions deter-
o
~ nn
•. c::7
UU
0
UW
....
c
nl
'------------------
,'- ." Vertebral cross-sections from autopsy specimens of
young (A) and old (8) bone show a marked reduction in
cancellous bone in the latter. Reprinted with permission
,I
from Nordin. B,E.e. (1973). Metabolic Bone and Stone Disease. Edinburgh: Churchill Livingstone. C. Bone reduction
I
l__
P
9 9
W_i'_h_a__i_n__i'_'_'_h_e_m_a_t_i,_a_'_'Y_d_e__i'_t_e_d_._A_,_n_o_,_m_a_1_b_o_n_e
mined by the geometry and n1mcrial properties of
each structure (Case Stud)' 2-2). A large plate, carrying high loads, unloads the bone to a great extent; the bone then atrophies in response to this diminished load, (The bone may hypertrophy at the
bone-screw interface in an altcmpt to reduce the
rnicrol1lotion of the screws.)
Bone resorption under a plate is illustrated in
Figure 2-48. A compression plate made of a material approximately 10 times stifTer than the bone \Vas
applied to a fractured ulna and remained after the
fracture had healed, The bone under the plate earried a lower land than normal; it was partially resorbed, and the diameter of the diaphysis became
markedly smaller. A reduction in the size of the
bone diameter greatly decreases bone strength. particularly in bending and torsion, as it reduces the
area and polar moments of inertia. A 20(M, decrease
in bone diameter may reduce the strength in torsion by 60%. Changes in bone size and shape illustrated in Figure 2-48 suggest that rigid plates
should be removed shortly after a fraeture has
healed and before the bone has markedly diminished in size. Such a decrea.sc in bone size is usually accompanied by secondary osteoporosis, which'
further weakens the bone (SI'itis et a!., 1980),
An implant may cause bone hypertrophy at its attachment sites. An example of bone hypertrophy
(top) is subjected to absorption (shaded area) durin
the aging process, the longitudinal trabeculae beco
thinner and some transverse trabeculae disappear
tom). Adapted from Sifter!, R.S .. & Levy. R.N. (J98/j. T
becular pacteffls and the internal architecture of bone.
S_i_"_"_i_J_rv_lo_d_,_";_.S_,_2_2_'_.
around scrcws is illustrated in Figure 2-49. A
plate was applied to a femoral neck fracture and
bone hypcrtrophied around the scrcws in resp
to the increased load at thesc sites. H.~lpert
may also result if bone is repeatcdly subjecte
high mechanical stresses within the normal ph
logical range. Hypertrophy of normal adult bon
response to strenuous exercise has been obse
(Dalen & Olsson, 1974; Huddleston et aI., 1
Jones et aI., 1977), as has an increase in bone
sity (Nilsson & Wesllin, 1971),
Degenerative Changes in Bone
Associated With Aging
A progressive loss of bone density has been
sCI·veci as part of the normal aging process. The
gitudinal trabeculae become thinner, and som
the transverse trabeculae arc resorbed (SifTe
Levy, 1981) (Fig. 2-50). The result is a marke
duction in the amount of cancellous bone a
thinning of conical bone. The relationship betw
bone mass, age. and gender is shown in Figure
The decrease ill bone Lissuc and the slight decr
in the size or the bone reduce bone strength
stillness.
ro
U
o
1000
E
w
'"
w
m
:;;
500
ID
C
o
tIl
20
~OWlng
40
60
BO
Age (years)
the relatIOnship between bone mass,
age, and gender. On the top of the figure, a crosssection of the diaphysis of the femur and the bone
mass configuration is shown. Reprinted ('lith permission
from Kaplan, F.S., rlayes, W.c., Keaveny, T.M., er al
(7994), Form and function of bone. In S.R. Simon (edJ
Orthopaedic Basic Science (.0, 767). Rosemont, IL:AAOS
Stress~strain curves for specimens from
adult tibiae of t\VO widely differing ages te
tension are shown in Figure 2~52. The u
stress was approximately the sarne for the
and the old bone. The old bone specimen
withstand only half the strain that the youn
could, indicating greater brittleness and a re
in energy storage capacit},. The reduction
density', strength, and stiffness results in in
bone fragility!. Age~related bone loss depen
number of factors, including genclel: age
menopause, endocrine abnormality, inactiv
use, and calcium deficiency. Over several d
the skeletal mass ma.y be reduced to 5000 of
trabecular and 25°/(; of cortical mass. In the
decade, women lose approximately 1.5 to 2(1
while men lose only approximately half th
(0.5 to 0.75(-/0) yearly. Regular physical activ
exercise (Zetterbarg et aI., 1990), calcium, a
sibly estrogen intake may decrease the rate
mineral loss during aging.
Summarv
1i Bone is a complex two-phase composit
rial. One phase is composed of inorganic
salts and the other is an organic matrix of c
and ground substance. The inorganic com
makes bone hard and rigid, whereas the
component gives bone its flexibility and resi
2 tVlicroscopically, the fundamental str
unit of bone is the osteon, or haversian
composed of concentric layers of a mineraliz
trix surrounding a central canal containin
vessels and nellie fibers.
Strain
Stress-strain curves for samples of adult young and
old human tibiae tested in tension, Note that the
bone strength is comparable but that the old bone is
more brittle and has lost its ability to deform.
Adapted from Burs rein, A.H., Reilly, D. T, & Alartens, M.
(976). Aging of bone tissue' Mechanical properties.
Bone Joint Surg, 58A, 82.
3· rv(acroscopicall.\" the skeleton is comp
cortical and cancellous (trabecular) bone.
bone has high density' while trabecular bon
in density over a wide range.
Bone is an anisotropic material, exhibi
ferent mechanical properties \vhen loaded in
ent directions. iVlature bone is strongest and
in compression.
S Bone is subjected to complex loading
during common physiological activities
walking and jogging. Most bone fractures
duced by a con1bination of several loading
p
s
a
m
6 Ivluscle contraction affects stress patt
bone by producing compressive stress that p
or totall).' neutralizcs thc tcnsilc stress acting on the
bone.
Bone is stiffer, sustains higher loads before failing, and stores more energy whcn loaded at higher
physiological strain rates.
~i Living bone fatigues when the frequcnc y.' of
loading precludes the remodeling necessary to prevent failure.
The mechanical behavior of a bone is influenced by its geometry (length, cross-sectional area,
and distribution of bone tissue around the neutral
Bone remodels in response to the mechanical
demands placed on it; it is laid down \vhere needed
and resorbed where not needed.
vVith aging comes a marked reduction in the
amount of cancellous bone and a decrease in the
thickness of cortical bone. These changes diminish
bone strength and stiffness.
REFERENCES
Bassett. C.A.L. (1965). Electrical effects in bone. Sci .-\111,2/3. 18.
BClIldicld, W.. 0.: Lt, e.H. (1967). AnisoLropy of nonelastic llow in
bone. J App{ Physics, 38, 2450.
Buckwaltcl~ lA.. Glimcher, \U., Coopel: R.R., et al. (1995). Bone biology. Part I: Stnlcture. blood supply, cells, matrix and mineralization. Pan II: Formation, form, remodelling and regubtion of
cdl funcLion. ((nstructional Course Lectlln:). J BOlle JoiJl! Sllrg.
77A, 1256-1289.
Burskin, A.H., Reilly, D.T.. 0.: \bnens, M. (1976). Aging of bone tissue: :vlechanical properties. J BOllI..' Joilll SI/1~!i. 58:\, 82
Burstein, A.H., L't al. (1972). Bone strength: The effecl of scre\\"
holes. J BOlle Joillt Slil~!i, 54:\, 1143.
n.R. (1978). Anisotropic analysis of strain roselle information from cortical bone. J BiolJlcch, II, 199.
D.R.. 0.: I-!ayes, W.e. (1977). Compact bone fatigue damage:
microscopic examination. Clill Of/hop, /27,265.
N.. 0.: Olsson, K.E. (1974). Bone mineral content .md physiactivity. Acla Orthop Scall(/, 45, 170.
G.U., et al. (1979). Bone densitometry using computed tomography. Part I: Selective determination of trabecular bone
and oLher bone mineral paraml'lers. Normal values in
children and adults. BrJ Radiol, 52, 14.
Frankel, VH., &: Burslein. A.H. (1970). Ort/lOpacdic Biol/li.'clul
Philadelphia: Lea 0.: Febigel:
f-Iuddbaoll, A.t.. Rockwell, D.. Kulund, D.N., et a1. (1980). B
mass in lifetime tennis athletes, hUlA, N4, 1107.
!nternaLional Societv of Biomechanics (1987). QII{/lIlitics alld l
oj" J1easllfi.'IIiCIIlS· iI/ Bio/lltxhal/ics (unpublished).
Jenkins. D.P.. 0.: Cochran, T.H. (1969). Osteoporosis: The dram
effecl of disuse of an extremity. Clill On/lOp, 64, 128.
Jones, H., Priest . ./.. Hayes, \V.. et a1. (1977). f-lu!11l.'ral hYPl.'nr
in response to exercise. J BOlle Joill/ SIiI;!5. 59:\, 204.
Kaplan. ES .. Hayes. We., Kea\'eny. T.iVI., l'l al. (1994). Form
funcLion of bone. In S.R. SiTllon (Ed.). Orlhopacdic Basic Sc
(pp.I27-184). Rosemont. IL: AAOS.
Kazarian, L.L., & \.lon Gil.'r'ke, I-I.E. (1969) Bone loss as a resu
immobilization and chelation. Preliminary ri:sults in :\1<
lllubtla. Clill Orthop. 65. 67.
Ke.weny, T.M .. &: Hayes, \V.e. (1993). Mechanical propertil.'s of
tical and trabecular bone. BOIIC, i, 285-344.
Krolner, B.. 0.: Toft. B. (1983). Vertebral bone loss: An unheeded
effecL of bedrest. C!ill Sci, 6e+. 537-540.
Kummer, J.K. (1999). Implant Biomatl:.'rials. In: L\1. Sph'ak,
DiCesare, D.S. Feldman, KJ. Ko\·al. A.S. Rokito. & J.D. Zu
man (Eds.). Or/hop(i('dics:"\ S/Ildy Gllidc (pp. 45-48). New
\IcGraw-I-lill.
Lanyon. LE.. &. Bourn. S. (1979). The inlluence or lllechanicd
Lion on the dc\'elopmcrH and remodeling of the tibia: An ex
mental study in sh('ep. .1 BOlle Joinl SI/rg. 6/.4. 263.
Lanyon, L.E., H.unpson. \\I.GJ .. Goodship, A.E., et al. (1975).
deformaLion recorded in vin) from sLrain g'lllgl.'S attached t
human tibial shaft. ..lew Onhop Sct/lul, "'6~ 256.
Nilsson. B.E., & \\btlin, N.E. (1971). Bone density in athletes.
On/lOp. 77, l79.
6zkaya, N., & Nordin, M. (1999). Fllndall/ell/tlls
Biolllcc!ul
EquilibriulII, ,HOlioll. alld Del(lnllatioll (2nd cd.) Ncw
Springer-Verlag.
Rarnballl~ P.c., 0.: ~Johnston. R.S. (1979). Prolonged weighLles
and calcium loss in man. Ac!a ASll'ollal/lica, 6, 1113.
Siffen, R.S .• & Levy, R.N. (1981). Tnlbccular patterns and the i
nal architecture of bonl.'. ;\fl. Sinai J lIed, 48. 221.
SltiLis, P., Paa\"{llainell, P., Karaharju, E.. ct al. (1980). Structura
biolllechanical changes in bOl~e aher rigid plate fixation. C
SlIIp., 23, 247.
\Vhedon, G.D. (1984') Disuse osteoporosis: Physiological asp
Cab!" 'fissile lilt, 36, 146-150.
Wolff, J. (1892). Dlls Gt:set::. der halls!(JrlII(ltiOiI der Kllochcll. B
I-lirschwald.
Zetterberg e., Nordin. :\'1., Sko\Ton. M.L.. et al. (1990). SkeleL
feClS of physical activity. Geri-7bpics. /3(4). 17-24.
or
C
VASCULAR FACTO,RS
J
I
MECHANICAL FACTORS
I
i
i'
SlruCturc
Function
,ilJ
;j"
w
Protection, suppOrt. kinCffiJtic links
~~ Organic Component
i
Collagen Proteins
1 ;,N~~coliag~n'Proteln~;
Type I )
~1
1
Growth factor cell
attachment proteins
PG',
\1
Inorganic Component
Cry,Olh of
HYDROXYAPATITE '
Calcium, Phosphate
1
1
:
.j
FLOW CHART 2·2
~This
flow chart is
dl?~igncd
Bone composition. structure, and functions.·
(PG's,
proteoglycans)
for classroom or group discussion. Flow chart is not meant to be exhaustive.
.1
'1
I
fJ, ,,1'.j2ij4if@_~\M¥i"I:IIi:i ;:H%J! dl
'I);; )'MltII: i)() 1tWC' !~
',~.
•~
a.
"
>
~
"a0
E
£
m
x
.."'"
.s
V
'"'m"
E
0
"
0
0
.0
B
cm
0
E
<;
c
m
.-
'"c0
{;
"'
u:
V
.0
em
~
0
.~
C
v
":;:
·u
'i3
0
0
m
2
E
m
~
~
§
u
.:"
u
.;;;
C
"5
x
w
0
~
0
'"5
E
0
~
m
u
.e
"
.,'"c
•
". 0
;;;
"'
~
0
'"
~
,.
]
I
j
·(11
Gcnedc DIsturbance
Blood Flow
<.;,\' :..\;.:, \,.::,;:.\-
,~":
Tumoral DIsturbance
"
J
.;',: 'i.~,:('erjm"ry
:j
,j
i
:\f~pvcloping'
\i ,;,:::';;'Skclc~on
1
\:<·",-':::::::Y",
I
J
6~'~'~6'~i';~ndro$is
/~\;~r~rth~s: N
,1
j
.Ostco~h'~n'd~i;i'~ ,
disseC30S Kcil:,
.Os,conccrosi~~~
"_fC~pr.'l1 head
~ 'Shcucrm3nn;
of the femoral j:,
,,,,:,)::,::~,'.i:;
·Gauchcrs
disease
·Cushing
Syndrome
condrlos'
\0"
, "';; .'
J"f:,Vcrccbrac'
IJ
\' .' ,,-.
,::
';:1
~,i
:j;
1/<1
.'~,
ri
PATHOLOGICAL OR FRAGILITY FRACTURE
ij
~
t~:
i~
o
I'"
FLOW CHART 2-4
Intrinsic factors associated with bone damage. Clinical examples.*
"This flow chart is designed for d<lssroom or group discussion. Flow chart is not meant to be cxhamtivc.
-''>-';;,i'
~C"' l'
:,i,¥
(;~
"
"-",
·,'i-".",,"';c.c
-'~,,;.;':.,
,:
Biomechanics of
Articular Cartilage
Van C. Mow, Clark T Hung
Introduction
Composition and Structure of Articular Cartilage
Collagen
Proteoglycan
WatE'r
Structural and PhysicallniE'raction Among Cartilage Components
Biomechanical Behavior of Articular Cartilage
Nature of Articular Cartilage Viscoelasticity
Confined Compression Explant Loading Configuration
Biphasic Creep Response of Articular Cartilage in Compression
Biphasic Stress-Relaxation ResponsE' of Articular Cartilage in
Compression
Permeability of Articular Cartilage
Behavior of Articular Cartilage Under Uniaxial Tension
Behavior of Articular Cartilage in Pure Shear
Swelling Behavior of Articular Cartilage
Lubrication of Articular Cartilage
Fluid-Film Lubrication
Boundary Lubrication
Mixed Lubrication
Role of Interstitial Fluid Pressurization in Joint Lubrication
Wear of Articular Cartilage
Hypotheses on Biomechanics of Cartilage Degeneration
Role of Biomechanical Factors
Implications on Chondrocyle Function
Summary
Acknowledgments
References
Flow Charts
Composition and Structure
Articular Cartilage
Introduction
Three types of joints exist in the human body: fib'ro~ls, cartilagin,olls. and synovial. Onl.y one of'
these,;-,ollle syn.ovial. or diarthrodial, jo.int, allows a
large degree of motion. In Y~_L1ng normal joints, the
,-~llticulalingJ)onecn~ls of diartl~~'odiaIj9_inls,!.re cov}--ered by a-thin (1-6 mill), den~c, tram~JJlcc_nt,~:-,'hile
,~~c6nnective-i.i.ssllc called hyaline arliCltlar. cartilage
iCBox-3:!). Articul'IU~f\rtilage-;"s a highly s-p·ecia·lized
. tissue precisely suited for withstanding the highly
i~adcd joint environment without failure during an
a\lerage individual's lifetime. PhXti.t~~l.Qgically, however, ,il is v51~~~I'.lIly an isolated li,ssuc, devQ.id of blood
\;essds, 1).~-mphalic channels, and nelyrol.ogical inncr~alion~'-Flll'lllermoJ'c,its .c~l.llliar density is less than
.·tK"t of any other tissue (Stoek\~c·il. 1979).
[n (liartll-i'o(Ti~;fioints, articular cartilage has two
- p-rimar'yfl~nctions': (1) to distl-ibute jo.int loads over
,a wiele al-ea, thus decreasing lhe strc:)ses sllst~!.!ned
by the contacting join.t..~.sl~rraccs (Atcshian ct aI.,
1995; Helminen et aI., 1987) and (2) to allow rel"tive
movementoft!1cu:m,posing joinl surfaces \~:·,i.t"I,l""I~lin­
imal ft:'iction and wear (iVlow & Alcshian, 1997). In
thi's"ch~l'I)iel','wc w'Hi describe how the biomcchanical pl'openics of articular cartilage, as determined
by its composition and structure, allow for the optimal performance or these functions.
~rticular
,
I
Cartilage
A notable exception to the definition of hyaline articular G'lrtilage is the temporomandibular joint, a synovial
joint in which fibrocartilage is found covering the bone
ends. Fibrocartilage and a third type of cartilage, elastic
cartilage. are closely related to hyaline cartilage embryologically and histologically but .'lre vastly different in
mechanical and biochemical properties. Fibrocartilage
represents a transitional cartilage found at the margins
of some joint cavities, in the joint capsules, and at the
insertions of ligaments and tendons into bone.
Fibrocartilage also forms the menisci interposed between the articular cartilage of some joints and composes the outer covering of the interverlebral discs. the
anulus fibrosus. Elastic cartilage is found in the external
ear, in the cartilage of the eustachian tube, in the
epiglouis. and in certain pans of the larynx.
~.~-
.,
of
Chondrocytcs, the sparsely distributed cells in ar
lllarc~\I:tiL\ge, account for less than 10% of the
sue's volume (Stoekwell. 1979). Schematically.
zonal arrangement of chondrocytcs is shown in
ure 3-1. Despite their sparse distribution, chond
cytcs manufacture, secrete, on!anil',---c, and rnain
tile on!anicco'n;-I;~I~l-~f the~ extracellular ma
(ECM) (Fos'lIlg & Hardingham, 1996; ·M~lir, 19
The org~nic ma~rix,is compo~~5:L~~[.:~,,(I(:nscp9.J~\
of fin:"-§?·I-(..1¥8~,,,(~'bril~ (n10stly t'ype-Jl~c-olh;~gen.,
miii6i' amounts .,;1' types V, VI, IX, and XI) that
enrneshcd..Jn. a cOnCC.tl.traled ~olu.~i().n. of proteo
cans (pcs) (l3ate;;'an et aI., I996;Eyre, 1980; M
-~1983)_ 1n nOITn,)1 arLiculaL_cartilage the colla
cOlllent ranges from 15 to 22% by wet weight
the PG con-tent fronl4to 7% bv
wet wei(Tht·
the
.'
:::::>
'
maining 60 to 85 % is walCI: inorganic salts.
small amounts of other matrix proteins, glycop
teins. and lipids (Mow & Ratcliffe, 1997). Colla
flbrils and PGs, each being capablc of form
structural nctworks of significant strength (Bro
& Silyn-Roberts, 1990; Kempson et aI., 19
Schmidt el aI., 1990; Zhu et aI., 1991, 1993), arc
structural components supporting the internal
chanical stresses that result rrom loads being
plied to the articular canilage, Moreover, th
structural components, together with water, de
mine the biomcchanical bchavior of this tis
(Ateshian et aI., 1997; Maroudas, 1979; Mow et
1980, 1984; Mow & Aleshian, 1997).
COLLAGEN
Collagen is the mosl abundant protein in the b
(Bateman et aI., 1996; Eyre, 1980). In articular ca
lage. collagen has a high level of structural organ
tion that providcs a fibrous ultrastructure (Cl
1985; Clarke, 1971; Mow & Ratcliffe, 1997). The
sic biological unit of collagen is tropocollagen
Sln.lcture composed of three procollagen polypep
chains (alpha chains) coiled into left-handed hel
(Fig. 3-2;\) thaI are furthcr coiled abollt each o
inlO a right-handed triple helix (Fig. 3-28). Th
rod-like tropocollagen molecules. 1.4 nanOJlle
(nm) in diameter and 300 nm long (Fig. 3-2, C &
polymerize into larger collagen fibrils (Batem
ct al.. 1996; Eyre, 1980), In articular cartilage, th
fibrils have an average' diamcter of 25 to 40
(Fig.3-2E, Box 3-2); howevec this is highly varia
......
Articular surface
STZ (10.20%)
,,,,'•.,,•.,0'
., ,,~) •..,.> _,~
.~ ,.,
," '.
Middle zone (40-60<::0)
'~,,~-' ,,-r:~ '~
,~.-r
,'-'
~{~:f2jJ.~
Deep zone (30%)
i~;~f
Calcified zone
";·'5c..C_~,.,,1::;\,,_, Subchondral bone
B
" Tide mark
Chondrocyte
Photomicrograph (A) and schematic representation (8) of the chondrocyte arrangement
throughout the depth of noncalcified articular cartilage. In the superficial tangential zone,
chondrocytes are oblong with their long axes aligned parallel to the articular surface. In
the middle zone, the chondrocytes are "round" and randomly distributed. Chondrocytes in
the deep zone are arranged in a columnar fashion oriented perpendicular to the tidemark,
the demarcation between the calcified and noncalcified tissue .
•
Scanning electron microscopic studies, for instance,
have described fibers with diameters ranging up to
200 nm (Clarke, 1971). Covalent cross-links form between these tropocollagen molecules, adding to the
flbrils high tensile strength (Bateman et aI., 1996).
The collagen in articular cartilage is inhomogeneously distributed, giving the tissue a layered character (Lane & Weiss, 1975; Mow & Ratc1iITc, 1997).
Numerous investigations Llsing light. transn1ission
electron, and scanning electron microscopy have
identifled three separate structural zones. For example, Mow et al. (1974) proposed a zonal arrangement for the collagen network sho\vn schematically'
in Figure 3-3A. In the superficial tangential zone,
which represents 10 to 20()0 of the total thickness,
are sheets of fine, densely packed fibers randomly
\voven in planes parallel to the articular surface
(Clarke, 1971; Redler & Zimny, 1970; Weiss et aI.,
1968). In the middle zone (40 to 60% of the total
thickness), there are greater distances between the
randomly oriented and homogeneousl::.. dispersed
fibers. Below this, in the deep zone (approximatel.v
30% of the total thickness), the fibers come together,
forming larger, radially oriented fiber bundles
(Redler et aI., 1975). These bundles then cross the
tidemark, the interface between articular cartilage
and the calcified cartilage beneath it, to enter the
calcified cartilage, thus forming an interlocking
"root" sj'stem anchoring the cartilage to the uncler-
lying bone (Bullough & Jagannath, 1983; Redl
aI., 1975). This anisotropic fiber orientation is
rored by the inhomogeneous zonal variations in
collagen content, which is highest at the surface
then remains relatively constant throughout
deeper zones (Lipshitz et aI., 1975). This comp
tionalla.vering appears to provide an important
mechanical function by' distributing the stress m
uniformly across the loaded regions
the join
sue (Selton et aI., 1995).
Cartilage is composed primarily of type II c
gen. [n addition, an array of difFerent coll
(t)'Pes V, VI. IX, Xl) can be found in quantitati
minor amounts within articular cartilage. Tvp
collagen is present primarily in articular cm·ti
the nasal septum, and sternal cartilage, as well
or
Differences in Coliagen Types
Differences in tropocollagen alpha chains in various
body tissues give rise to specific molecular species, or
types of collagen. The collagen type in hyaline cartilage
type II collagen, differs from type I collagen found in
bone, ligament, and tendon. Type II collagen forms a
thinner fibril than that of type I, permitting maximum
dispersion of collagen throughout the cartilage tissue.
(r
A
I
l
1.4 nm
B
chain
Triple helix
Tropocollagen
molecule
o
Collagen fibril with
quarter-stagger array
of molecules
Fibril with repeated
banding pattern seen
under eleclron microscope
Molecular features of collagen structure from the alpha chain ((I) to the fibril. The flexible
amino acid sequence in the alpha chain (A) allows these chains to wind tightly into a righthanded triple helix configuration (8), thus forming the tropocollagen molecule (C). This
tight triple helical arrangement of the chains contributes to the high tensile strength of
the collagen fibril. The parallel alignment of the individual tropocollagen molecules, in
which each molecule overlaps the other by about one quarter of its length (0), results in a
repeating banded pattern of the collagen fibril seen by eledron microscopy (x20,OOO) (E).
Reprinted wirh permission from Donohue, It'l1., Buss. D., Oegema. IR., et al. (19831- The effects of
indireCl blunt trauma Oil adult canine arricuhu cartilage. J Bone Joint Surg, GSA. 948.
•
the inner regions of the intervertebral disc and
meniscus. For reference, type I is the most abundant collagen in the human body and can be found
in bone and soft tissues such as intervertebral discs
(rnainl.y in the annulus fibrosis), skin. meniscus, tendons, and ligaments. The most important mechanical properties of collagen fibers nrc their tensile
stiffness and their strength (Fig. 3-4;\). Although a
Single collagen fibril has'-not be~n tested in len;ion.
the tensile slI'ength of collagen can be inferred from
tests on structures with high collagen contenl. Tendons, for example, are about 80% collagen (dry
weight) and have a tensile stiffness of 10.1 !\'1Pa and
a tensile strength of SO MPa (Akizuki et aI., 1986;
Kempson, 1976, 1979; Wooet aI., 1987, 1997).Stec1,
by comparison, has a tensile stillness of approximately 220 x 10" MPa. Although strong in tension.
collagen fibrils offer little resis(ance lo ~olllpression
because their large slendcnlcss ratio, the rati
length to thickness, makes it easy for them to bu
under cotl1prcssive loads (Fig. 3-48).
Like bone, articular canilage is anisotropic
material properties differ with the direction of l
ing (Akizuki et aI., 1986; Kempson, 1979; Mow
RalclifTc, 1997; Roth & Mow, 1980; Woo el
1987). Oft is thought that this anisotropy is relate
the varying collagen fiber arrangements within
planes parallel to the articular surface. It is
thought, however, that variations in collagen f
cross~link density, as well as variations in colla
PG interactions. also contribute to articular c
lage tensile anisotrop~", In tension, this anisotr
is llsually described with respect to the direc
of the articular surface split lines. These split l
arc elongated fissures Jj·roduced by piercing the
ticular surface with a small round awl (Fig.
B
STL
A. Schematic representation, (Repriflleej \·virh permission from
Mow v.c. el al. j1974j. Some surface dJc1riJcteristics of arriculaf
cartilages. A scanning electron microscopy sludy and a theoretical
mode! for the dynamic interaction of synovial fluid and articular car·
tilage. J Bionll?{hanics, 7, 449), B. Photomicrographs (x3000;
provided through the courtesy of Dr. T. Takei, Nagano. Japan)
of the ultrastructural arrangement of the collagen network
throughout the depth of articular cartilage. In the superficial
tangential zone (STZ), collagen fibrils are tightly woven into
sheets arranged parallel to the articular surface. In the middle
Middle zone
Deep zone
zone, randomly arrayed fibrils are less densely packed to accommodate the high concentration of proteoglycans and water, The collagen fibrils of the deep zone form larger radially
oriented fiber bundles that cross the tidemark, enter the calci·
fied zone, and anchor the tissue to the underlying bone. Note
the correspondence between this collagen fiber architecture
and the spatial arrangement of the chondrocytes shown in
Figure 3-1. In the above photomicrographs (B), the STZ is
shown under compressive loading while the middle and deep
zones are unloaded.
Collagen Fibril
<)0
A
B
1m))
))))))
1m))
)l»ll
IllJli
lJIlil I (>
High tensile sliflness and strength
Uttle resistance to compression
Human femoral condyles
Illustration of the mechanical properties of collagen fibrils:
(A) stiff and strong in tension, but (B) weak and buckling
easily with compression. Ad,lpted from Myers, E.R., Lai, WM.,
& MOV'I, VC (1984). A COfltinuum theory and all experiment for
the ion-induced swelling be/Mvior carlilage. j Biomech Eng,
Gelenkknorpel. Verhandlungen der Analomischen Gesellschaft,
106(21. 15/-/58.
12.248.
Diagrammatic representation of a split line pattern on the
surface of human femoral condyles. Reprinted wirh permission from Hu/rkrantz, W (1898). Ueber die Spafrrichwflgefl der
Hultkrantz, 1898). The origin of the pattcrn is related to the directional variation 01" the tensile stilTness and strength characteristics of articular can ilage described above. To dale. ho\Vevel~ the exact
rc;sons as to why articular canilage exhibits slIch
pronounced anisolropies in tension is not known,
nor is the functional significance
anisotrop)'.
or
this tensile
PROTEOGLYCAN
l\llany types of PGs arc round in cartilage. Fundamentally, it is a large protein-polysaccharide mole-
cule composed of a protein core to which one or
more glycosaminoglycans (GAGs) are attached
(Fosang & Hardingham, 1996; Muir, 1983; Ratcliffe
& Mow, 1996). Even the smallest of these molecules,
high'Can and deem-in, are quite large (approxim;t~ly I X 10'\ 111\\'). but they comprise less than
10~~ of all PGs present in the tissue. Aggrccans are
much larger (1-4 X 10" mw), and they have the rcmarkable capability to attach to a hyaluronan mol·
ccule (HA: 5 X 10' I11W) via a specific 1·-1A-binding rcgion (HABR). This binding is stabilized by a link
protein (LP) (40-48 X 10" mw). Stabilization is crucial 10 the function of normal cartilage; without it,
the components of the PG molecule would rapidly
escape from the tissue (Hardingham & IVluir, 1974;
H<lscall, 1977; Muil; 1983).
Two types of GAGs comprise aggrecan: chondroitin sulfatc (CS) and kcmtan sulfate (KS). Each
CS chain contains 25 to 30 disaccharide units, while
the shorter KS chain contains 13 disaccharide units
(MuiJ~ 1983). Aggrecans (previously rererred to as
subunits in the American literature or as monomers
in the UK and European literature) consist of an approximately 200.nanol11ctcr..long protein core to
which approximately 150 GAG chains, and both 0linked and N-linked oligosaccharides, are covalently
attached (Fosang & Hardingham, 1996; Muir, 1983).
Furthermore, the distribution of GAGs along the
protein core is heterogeneous; there is a region rich
in KS and O-linked oligosaccharides and a rcgion
rich in CS (Fig. 3-6M. Figure 3-611 dcpicts the famous "boule-brush" model for an aggrecan (j\lluir,
1983j. Also shown in Figurc 3-6/\ is the hctcrogcneity of the protein core that contains three globular
regions: GIl the HABR located at the N-terminus that
contains a small amount of KS (Poole, 1986) and a
few N-Iinked oligosaccharidcs, G., located between
the HABR- and the KS-rich regio·n (Hardingham et
aI., 1987), and G" the corc protcin C-tenninus. A I: I
stoichiomelly C:dslS between the LP and the GI
binding region in c.:anilagc. iVlore recently, the o
two globular regions have been extensively stu
(Fosang & Hanlingham. 1996), but their functio
significance has notycl been elucidated. Figure 3
is lhe accepted molecular confonnation of a PG
gregate; Rosenberg Gt al. (1975) were the firs
obtain an electron micrograph 01" this mole
(Fig. 3-6C).
In native cartilage, most aggrccans arc associ
withHA to form the large PG aggregates (Fig. 3These aggregates may have up to several hund
aggrecans noncovalenliy uuached to a central
core via their HABR, and each sile is stabilized
an LP. The filamentous HA core molecule is a n
sulfated disaccharide chain that may be as long
fJ.m ill Icngth. PG biochcmists have dubbed thc
an "honorary" PG, as it is so intimately involve
the structure of the PG aggregate in articular c
lage. The stability afforded by the PG aggregates
a rl"wjor functional significance. It is accepted
thut PG aggregation promotes immobilization
the PGs within the nne collagen meshwork, add
structural stability and rigidity to thc ECM (Mo
aI., 1989b; Muir, 1983; Ratcliffc et aI., 1986).
thermore, two additional forms of c!crmatan sul
PG have been idcntif"led in the ECM of articular
tilage (Rosenberg et aI., 1985). In tendon, c1erm
sulfate PGs have been shown to bind noncovale
to the surfaces of collagen fibrils (Scott & Orf
1981); however, the role of derma tan sulfatc in
ticular cartilage is unknown, biologically and f
tionally.
Although aggrecans generally have the b
structure as described abovc, they arc not st
turally identical (Fosang & Hardingham, 1996).
grecans vary in length. molecular weight, and c
position in a variety of ways; in other words,
arc polydisperse. Studies have demonstrated
distinct populations of aggrecans (Buckwalter e
1985; Heincgard ct aI., 1985). The first populatio
present throughout life and is rich in CS; the sec
contains PGs rich in KS and is present only in a
cartilage. As articular cartilage matures, other
related changes in PG composition and struc
occur. vVith canilage maluration, the watcr con
(Armstrong & Mow, 1982; Bollet & Nancc, 1
Linn & Sokoloff, 1965; Maroudas, 1979; V
1978) and the carbohydrate/protein ratio prog
sively decrease (Garg & Swann, 1981; Roughle
White, 1980). This dccrcase is mirrored by a
crease in the CS content. Conversely, KS. whic
present only in small ,imounts at birth, incre
throughollt development and aging. Thus,
Chondroitin Sulfate
(CS) chains
C-terminal
globular
domain (G 3)
A
Proteoglycan (PG) Macromolecule
1200 nm
Hyaluronan (HA)
B
1DI1I'----
A. Schematic depiction of aggrecan, which is compose
keratan sulfate and chondroitin sulfate chains bound c
lently to a protein core molecule. The proteoglycan pr
core has three globular regions as well as keratan sulfa
rich and chondroitin sulfate-rich regions. B. Schematic
resentation of a proteoglycan macromolecule. In the m
trix, aggrecan non covalently binds to HA to form a
macromolecule with a molecular weight of approxima
200:<. 106. Link protein stabilizes this interaction betw
the binding region of the aggrecan and the HA core m
cule. C. Dark field electron micrograph of a proteoglyc
aggregate from bovine humeral articular cartilage
(>:120,000). Horizontal fine at lower right represents O
Reprinted with permission from Rosenberg, C, Hellmann, W
Klein5chmidr. AX (975). Electron microscopic studies 0; p
gfycan aggregates (rom bovine articular cartifage. J Bioi Che
250. 1877.
i
If
II!!!..".""
~", ~ ", ~""~"::-'~~·,.cT·","'! '.," '.,=-, ~ ----:'....,....-----~.
,IIIIIIIIII!!I,II,'IIIIlIl!!.,!,!!!,,.p.
•.
"!!"!'!.
••••
.. "'-.
CS/KS ratio, which is approximately 10: 1 at birth, is
only approxirnatel:v 2: I in adult cartilage (Roughlcy
& "Vhile, 1980: Sweet et al" 1979; Thonar et al"
1986). Furthermore, sulfation of the CS molecules,
\vhich can occlir at either the 6 or the 4 position,
also undergoes age-related changes. In utero, chondroitin-6-sulfate and chondroitin-4-sulfate are present in equal molar amounts; however, by maturity,
the chondroitin-6-sulfate:chonclroitin-4-sulfatc ratio has increaseel to approximately 25: I (Roughley
et a1., 1981). Other studies have also documented an
age-related decrease in the hydrodynamic size of
tf~e aggrecan. Many' of these early changes seen in
articular cartilage may reflect cartilage maturation,
possibly as a result of increased functional demand
\vith increased \vcight-bcaring. Ho\vever, the functional significance of these changes, as well as those
occurring later in life, is as .\'ct undetermined.
WATER
\Valel~ the most abundant component of articular
cartilage, is n10st concentrated ncar the anicular
surface (_80°;(j) and decreases in a near-linear fashion with increasing depth to a concentration of approximately 65% in the deep zone (Lipshitz et al.,
1976: Maroudas, 1979). This Iluid contains many
free mobile cations (e.g., Nu', K', and Cal.) that
greatly! influence the mechanical and physicochemical behaviors of cartilage (Gu et aI., 1998; Lai et aI.,
1991; Linn & Sokoloff, 1965; Maroudas, 1979). The
fluid component of articular cartilage is also essential to the health of this avascular tissue because it
permits gas, nutrient, and waste product movement
back and forth between chondrocytes and the surrounding nutrient-rich synovial fluid (Bollet &
Nance, 1965; Linn & Sokoloff, 1965; Mankin &
Thrashel; 1975; Maroudas, 1975, 1979).
A small percentage of the \vater in cartilage resides intracellularl y', and approximately 30% is
strongly associated with the collagen fibrils
(Maroudas et aI., 1991; Torzilli et aI., 1982). The interaction between collagen, PG, and water, via Donnan osmotic pressure, is believcd to have an important function in regulating the structural
organization of the ECM and its s\velling properties
(Donnan, 1924; Maroudas, 1968, 1975). Most of the
\vater thus occupies the interfIbrillar space of the
ECM and is free to move whcn a load or pressure
gradient or other electrochemical motive forces are
applied to the tissue (Gu et aI., 1998; Maroudas,
1979). When loaded by a compressive force, ap-
proximately 70(;r; of the water may be moved, T
interstitial fluid movement is important in cont
ling cartilage mechanical behavior and joint lu
cation (Ateshian et aI., 1997, 1998; Hlavacek, 1
Hou et aI., 1992; Mow et aI., 1980; Mow & Atesh
1997).
STRUCTURAL AND PHYSICAL INTERACTION
AMONG CARTILAGE COMPONENTS
The chemical structure and physical interacti
of the PG aggregates influence the properties of
ECM (Ratcliffe & Mow, 1996). The closely spa
(5-15 angstroms) sulfate and carboxyl cha
groups on the CS and KS chains dissociate in s
tion at physiological pH (Fig. 3-7), leaving a h
concentration of fixed negative charges that cre
strong intramolecular and intermolecular cha
charge repulsive forces; the colligative sum
these forces (when the tissue is immerscd i
physiological saline solution) is equivalent to
Donnan osmotic pressure (Buschmann & Grod
sky, 1995; Donnan, 1924; Gu et aI., 1998; Lai et
1991). Structurally, these charge-charge repul
forces tend to extend and stiffen the PG ma
molecules into the interfibrillar space formed
the surrounding collagen network, To apprcc
the magnitude of this force, according to Step
Hawkings (1988), this electrical repulsion is
million, million, million, million, million, mill
million times (42 zeros) greater than gravitatio
forces.
In nature, a charged body cannot persist l
without discharging or attracting counter-ion
maintain electroneutrality, Thus, the charged sul
and carboxyl groups flxed along the PGs in artic
cartilage must attract various counter-ions and
ions (mainly Na', Cal., and C I") into the tissu
maintain electroneutrality. The total concentra
of these counter-ions and co-ions is given by
well-known Donnan equilibrium ion distribu
law (Donnan, 1924). Inside the tissue, the mo
counter-ions and co-ions form a cloud surround
the fixed sulfate and carboxyl charges, thus shield
these charges from each othet: This charge shield
acts to diminish the very large electrical repul
forces that otherwisc would exist. The net result
swelling pressure given by the Donnan osmotic p
sure law (Buschmann & Grodzinsky, 1995; Don
1924; Gu et aI., 1998; Lai et aI., 1991; Schuber
Hamerman, 1968). The. Donnan osmotic pres
theory' has been extensivel:v used to calculate
Aggregate Domain
Repulsive forces due 10 fixed
charge density distribution
A
Smaller domain
Increased charge density
Increased charge-charge repulsive forces
B
A, Schematic representation of a proteoglycan aggregate
solution domain (left) and the repelling forces associated
with the fixed negative charge groups on the GAGs of aggrecan (right). These repulsive forces cause the aggregate
to assume a stiHly extended conformation, occupying a
large solution domain, B. Applied compressive stress decreases the aggregate solution domain (left), which in turn
increases the charge density and thus the intermolecular
charge repulsive forces (right).
swelling pressures of anicular cartilage and the
intervertebral disc (Maroudas, 1979; Urban &
McMullin, 1985), By Starling's law, this swclling
pressure is, in llll'n. resisted and balanced by tension
developed in the collagen network, confining the
PCs to only 20% of their fTee Solulion domain
(Maroudas, 1976; Mow & Ratcliffe. 1997; Sellow
eL aI., 1995). Consequently. this swelling pressure
subjects the collagen network to a "pre-stress" of sig·
nificant rnagnitudL: c\'en in the absclH.':c or cx
loads (Sellon cl aI., 1995, 1998).
Canilage PGs arc inhomogeneollsl.\' dislri
th,'oughoul the malrix, with their concenlralion
erally being highest in the middle zone and low
the superficial and deep zOlles (Lipshitz Cl aI.,
,'Jlaroudas, 1968, 1979; Vcnn, 1978). Thc biome
ical consequence or this inhomogeneous swelli
havior of cartilage (caused b.v the varying PG co
throughout the depth of the tissue) has recently
quantitalively assessed (Sellon el al., 1998), Al
sults from n:cent finite clement calculations
on models incorporating an inhomogeneous P
lribution show lhat it has a proround effect o
interstitial counler-ion dislribulion throughou
depth of the lissuc (Sun et aI., 1998).
\·Vhen a compressive stress is applied lo th
lilage surface. there is an inSlantaneous defo
lion caused primal"i)~' by a change in the PG
ecular domain, Figure 3-7 B. This cxtc.'rnal
causes the inlernal prcsstll"e in the matrix
ceed the swelling pressure and thus liquid w
gin to flow out of" tht: tissue. As the fluid r!()\V
the PG concentration increases, which in tu
creases the Donnan osmotic swelling pressu
the charge-charge repulsive force and bulk
pressive stress unlil they are in equilibrium
the external stress. In this manner, the phy
chcmical properties of the PC gcl lrapped w
the collagen network enable il to resisl com
sion. This mechanism complements the
pla~:ed by collagen that, as previollsly desc
is strong in lC'nsion but wcnk in compre
The abilit.'· of PCs to resist compression
arises fl'om two sources: (I) the Donnan os
swelling pressure associated with the t
packed fixed anionic groups on the GAGS
(2) the bulk eomprcssivc stillness of the coll
PG solid matrix, Experimentally, the Donna
motic pressure ranges from 0.05 to 0.35
(Maroudas, 1979), while the elastic modulus
collagen-PG solid matrix ranges from 0.5
MPa (Armstrong & Mow. 1982; Alhanasiou
1991; Mow & Ratcliffe, 1997).
Il is now apparenl that collagen and PG
interact and thal these interactions m'e of
functional imponance. A small portion or th
have been shown to be closely associated wit
lagen and may serve as a bonding agent be
lhe collagen fibrils. spanning distnnccs thL
too great for collagen cJ."oss·links to develop (Ba
~t al.. 1996; Mow & Ratcliffe, 1997; Muir,
PGs arc also thought to play an important role in
maintaining the ordered structure and mechanical properties of the collagen fibrils (Muir, 1983;
Scott & Orford, 1981). Recent investigations show
that in concentrated solutions, PGs interact with
each other to form networks of significant
strength (Mow et aI., 1989b; Zhu et aI., 1991,
.i 996)..Morcover, the density and strength of the
interaction sites forming the network \vere shown
to depend on the presence of LP between aggre~
cans and aggregates, as well as collagen. Evidence
sl.w:gests that there are fewer aggregates, and
m(;~e biglycans and decorins than aggrecans, in
the superficial zone of articular cartilage. Thus,
there mllst be a difference in the interaction bel\VCCn these PGs and the collagen fibrils from the
superficial zone than from those of the deeper
zones (Poole et aI., 1986). Indeed, the inle"action
bet\veen PG and collagen not only' plays a direct
role in the organization of the ECr.,ll but also contributes directly' to the mechanical properties of
the tissue (Kempson el aI., 1976; Schmidt el aL.
1990; Zhu el aI., 1993).
The specific characteristics of the physical.
chemical. and mechanical interactions between
coHagen and PG have not yet been fully' deter·
mi.ned. Nevertheless, as discussed above, we know
that these structural macromolecules interact to
form a porous·permeable, fiber-reinforced composite matrix possessing all the essential mechanical characteristics of a solid that is swollen with
\vat.er and ions and that is able to resist the high
stresses and strains of joint articulation (Andriacchi
etal .. 1997; I-lodge el aI., 1986; Mow & Ateshian,
1997; Paul, 1976). It has been demonstraled lhat
these collagen-PG interactions involve an aggrecall, an I-IA filan1ent, type II collagen, other minor
collagen types, an unknown bonding agent, and
possibly smaller cartilage components such as collagen type IX, recently identified glycoproteins,
and/or polymeric HA (Poole ct aI., 1986). A
schematic diagram depicting the structural
arrangement \vithin a small volume of articular
cartilage is shown in Figure 3-8.
\Vhen articular cartilage is subjected to external
loads, the collagen-PG solid matrix and interstitial
lIuid function together in a unique way to protect
against high levels of stress and strain developing in the ECM. Furthermore, changes to the biochemical composition and structur,,;l organization
of lhc ECM, such as during osteoarthritis (OA), are
paralleled b:y changes to t';e biomechanical proper-
Hyalur
n
-'I,-\-----Interstltlal flU
Collagen fi
Schematic representation of the molecular organizatio
cartilage, The structural components of cartilage, colla
and proteoglycans, interact to form a porous composit
fiber-reinforced organic solid matrix that is swollen wit
water. Aggrecans bind covalently to HA to form large
teoglycan macromolecules.
ties of cartilage. In the following section, the
havior of articular cartilage under loading and
mechanisms of cartilage fluid flow will be discus
in detail.
Biol71echanical Behavior o{
Articular Cartilage
The biomechanical behavior of articular carti
can best be understood when the tissue is viewe
a multiphasic medium. In the present context, a
ular cartilage will be treated as biphasic mate
consisting of two intrinsically incompressible,
miscible, and distinct phases (Bachrach et aI., 1
Mow et aI., 1980): an inlerstitial fluid phase an
porous-permeable solid phase (i.e., the ECM).
explicit analysis 01' the contribution 01' the
charges and ions, one would have to consider t
distinct phases: a fluid phase, an ion phase, an
charged solid phase (Gu el al .. 1998; Lai et aI., 19
For understanding how the water contributes t
mechanical propcnies, in the prc.:SCnl context. arti<,:.
lIlar cartilage may be considered as a Ouid-filled
porous-permeable (uncharged) biphasic medium,
with each constituent playing a role in the functional
behavior of cartilage.
During joint arlicuhllion, forces at the joint surface may vary" from alnl0S1 zero to more than ten
times body weight (Alldriacchi et ai., 1997; Paul,
1976). The contact arcas also vary in a cOll1plcx
manner and typically they arc only or the order of
several square centimeters (Ahmed & Burke. 1983;
Ateshian et al .. 1994). It is estimated that the peak
contact stress may reach 20 MPa in the hip while
rising from a chair and to IVIPa during stair climbing (Hodge et aI., 1986; Newberrv et aI., 1997). Thus,
articular cartilage, under physiologicallonding conditions, is a highly stressed malcrial. To understand
how this tissue responds LInder these high physiological loading conditions. its intl"insic mechanical
properties in compression, tension, and shear must
be determined. From these properties, one can un·
dcrstand the load-can~ying mechanisms within the
ECM. Accordingly, the following subsections \\"ill
characterize tht: tisslIC' behavior under these loading
Il'lOclalities.
NATURE OF ARTICULAR CARTILAGE
VISCOELASTICITY
If a material is subjected to the action of a constant
(time-independent) load or a constant deforrnation
and its response varies with time, then the mechanical behavior of the material is said to be viscoelastic. (n general, the response of slich a material can
be theoretically modeled as a cornbination of the response of a viscolls Ouid (dashpot) and an elastic
solid (spring), hence viscoelastic.
The two fundamental responses of a viscoelaslic
material arc creep and stress relaxation, Creep occurs when a viscoelastic solid is subjected to the action of a constant load, Typically. a viscoelastic solid
responds with a rapid initial deforrnation followed
by a slow (time-dependent), progressivel)' increasing deformation known as creep until an equilibrium state is reached. Stress relaxation occurs when
a viscoelastic solid is subjected to the action of a
constant deformation, l~ypically. a dscoelastic solid
responds with a rapid, high initial stress followed by
a slow (time~dependent). progressively decreasing
stress required to maintain the deformation; this
phenomenon is known as stress relaxation.
Creep and strc~s rL,lnxalion phenomcna ll1a~' be
caused b~' differenl Illcchani:sms, For single-phase
solid pol~'lTh:ric materials, these phenomena are the
result of iJHcrnal friction caused b.\' the mOl ion or the
long polymL'ric chains sliding over ('ach olher \\'ithin
the stressed lll~lIcrial (Fling, 1981), The viscocl"stic
bdmvior of tendons and ligamt::IHs is primarily
eauscd bv this mechanism (Woo et aI., 1987, 1997).
For bone, the long-term viscoelastic lK'hador is
thought to ilL' caused h.\' a rdati\'c slip uf lamellae
within the osteons along with the Ilow of the inlL'rsti
tial Iluid (Lakcs & Saha, 1979). For articul,,,' cartihlgC. the compressive \'iscoelastic bdHl\'ior is prilllar~
i1~' caused h~' the Jlo\\' of tilL' interstitial lillie! and lhe
frictional drag associated with this lIow (Ateshian cl
aI., 1997; Mow (,t aI., 1980, 1984). In shear, as in
single-phase viscoelastic pol.vmers. it is primarily
c<.lusL·d b~' tilt:' motion of 10l1g pol~!J1lL'r chains such as
collagen and PGs (Zhu d aI., 1993, 1996). Thc component or anicu[;.\r cartilage dscoclasticit~,caused by
interstitial lluid !low is known as the hi phasic viscot.,lastic behavior (i'VlO\\' et a!., 1980), ~lI1d the COI11ponent of dscoebslicil~' caused b~' macromolecular
motion is known as til(' l1ow-indepl'ndelll (Ha,\'cs
Bc)dinc, 1(78) or the intrinsic visc()elastic behavior
Ihe collagen-PG solid matrix.
Although the deformational bdwyior has been
described in terms of n lillem' elastic solid (Hirsch,
1944) or viscoclastic solid (I-laves & Mockros, 1971),
lhes~ modcls fail to recognii'.e the role or water in
the \'iscoelastic behavior or and thc significant contribution lilal lIuid pressurizalion pla.vs in joint load
support and canilage lubrication (Ateshian L't aI.,
1998; Elmol'c et aI., 1963, Mow & Ratcliffe, 1997;
SokolofL 1963). Recentlv, experimental measurements ha\'c deh:nnined that interstitial fluid pressurization supports mort' than YOOk of the applied
load to the canilage surface (Solt/. & Ateshian.
1998) immediately following loading. This ('lfcC
can persist for more than 1,000 scconds and thus
shields til(.' EC1VI and chondroc.\'tcs from the crllshing defonnations or the high stresses (20 MPa) resulting frollljoilll loading.
'*
or
CONFINED COMPRESSION EXPLANT
LOADING CONFIGURATION
The loading of cartilage in vivo is cxtr('mel~" complex. To achic\'C' a better understanding of the deformational behavior or the tissue under load, an
explant loading conliguration known as confined
compression (lvlow et aL, 1980) has been adopted by
researchers. In this configuration, a cylindrical cartilage specimen is fitted snugly il1to a cylindrical,
srnZoth-walled (ideally frictionless) confining ring
that prohibits motion and fluid loss in the radial direction. Under an axial loading condition via a rigid
p~rol1s-permeable loading platen (Fig. 3-9;1), fluid
will flo\v from the tissue into the porous-permeable
platen, and, as this occurs, ~he cartilage samp~e will
compress in creep_ At any tIme the amount 01 compression equals the volume of nuid loss because
both the watcr and theECM arc each intrinsicall:v
incompressible (Bachrach et aI., 1998). The advantage of the confined compression tcst is that it creates a uniaxial, one-dimensional flow and def()rma~
tional J1eld within the tissue, which does not depend
on tissue anisotropy.' or properties in the radial diN
recLion. This greatly simplir-Ies the mathematics
needed to solve the problem.
It should be emphasized that the stress-strain,
pressure, fluid, and ion flow fields generated within
the tissue during loading can only be calculated;
ho\VevcI~ these calculations are of idealized models
and testing conditions. There are man:\' confounding factors, such as the time~dependent nature and
magnitude of loading and alterations in the natural
state of prc~strcss (acting \vithin the tissue), that
arise from disruption of the collagen network dur~
ing specimen harvesting. Despite limitations in de~
terrnining the natural physiological states of stress
and strain within the tissue in vivo, a number of re
searchers have made gains to\\'ard an understand~
ing of potential mechanosignal transduction I11ech~
anisms in cartilage through the use of explant
loading studies (Bachrach et aI., 1995; Buschmann
et aI., 1992; Kim et aI., 1994; Valhmu et aI., 1998)
based on the biphasic constitutive law For soft hy'drated tissues (Mow et aI., 1980).
N
BlPHAS1C CREEP RESPONSE OF ARTICULAR
CARTILAGE IN COMPRESSION
The biphasic creep response of articular cartilage
in a oneNdimensional confined compression experiment is depicted in Figure 3-9. In this case, a
constant compressive stress (To) is applied to the
tissue at tinlC to (point A in Fig. 3-98) and the tis~
Slie is allo\ved to creep to its final equilibrium
strain (EX). For articular cartilage, as illustrated
in the top diagrams, creep is caused by the exuN
dation of the interstitial fluid. Exudation is most
rapid initially', as evidenced by the early rapid r
of increased deformation, and it diminishes gra
ually until flow cessation occurs. During cre
the load applied at the surface is balanced by
compressive stress developed within the collag
PG solid matrix and the frictional drag genera
by the flow of the interstitial fluid during exu
tion. Creep ceases when the compressive str
developed within the solid matrix is sufficient
balance the applied stress alone; at this point
fluid flows and the equilibrium strain EX
reached.
Typically, for relatively thick human and bov
articular cartilages, 2 to 4 mm, it takes 4 to
hours to reach creep equilibrium. For rabbit c
tilage, which is generally less than 1.0 111m thi
it takes approximately' I hour to reach creep eq
librium. Theoretically, it can be shown that
time it takes to reach creep equilibrium var
inversely with the square or the thickness of
tissue (Mow et aI., 1980). Under relatively h
loading conditions, > 1.0 J\ilPa, 50 0/() of the to
fluid content may be squeezed From the tiss
(Echvards, 1967). Furthermore, in vitro stud
demonstrate that if the tissue is immersed
physiological saline, this exuded fluid is fully
coverable when the load is removed (Elmore
aI., 1963; Sokoloff, 1963).
Because the rate of creep is governed b.y the r
of fluid exudation, it can be used to determine
permeability coefficient of the tissue (Mow et
1980, 1989a). This is known as the indirect m
surement for tissue permeability (k). Average val
of normal hun1an, bovine, and canine pate
groove articular cartilage permeability k obtained
this manner arc 2.17 X 10. 15 M·'/N·s, 1.42 x 10'" M'/N
and 0.9342 x 10. 15 M4/N·s, respectively (Athanas
et aI., 1991). At equilibrium, no fluid flow occ
and thus the equilibrium deformation can be u
to measure the intrinsic compressive modulus (H
01' the collagen-PG solid matrix (Armstrong & M
1982; Mow et aI., 1980). Average values of norm
human, bovine, and canine patellar groove articu
cartilage compressive modulus H,.\ are 0.53, 0.
and 0.55 megapascal (MPa; note 1.0 MPa =
Ib/in 2 ), respectively. Because these coefficients ar
measure of the intrinsic material properties of
solid matrix, it is therefore meaningful to determ
ho\v they vary \vith matrix composition. It was
termined that k varies directly, while He' varies
versely with water content and varies directly w
PG content (Mow & Ratcliffe, 1997),
Confining ring
Impermeable platen
A
Unloaded (A)
~
oJ
"0
w
>
1
o
a
.~ E-'
c
.~
r
Equilibrium (C)
Creep (6)
E
a
0;
B
o
U_A_·___•__•
:e>
0-
w
•
•__•
Time
B
A. A schematic of the confined compression loading configuration. A cylindrical tissue specimen is positioned tightly into
an impermeable confining ring that does not permit deformation (or fluid flow) in the radial direction. Under loading,
fluid exudation occurs through the porous platen in the vertical direction. B. A constant stress (T" applied to a sample of
articular cartilage (bottom left) and creep response of the
sample under the constant applied stress (bottom right). The
o
No exudation C
- - =-.........._----...-
i
6
Copious
fluid
exudation
Time
Equilibrium
deformation
1
drawings of a block of tissue above the curves illustra
creep is accompanied by copious exudation of fluid fr
sample and that the rate of exudation decreases over
from points A to B to C. At equilibrium (EO'-'), fluid flow
and the load is borne entirely by the solid matrix (poi
Adapted from MOH~ VC, Kuei, S,C, Lai, Wl'vl., f:.'( af. (1980
Biphasic creep and stress relaxation of articular cartilage in
pression: Theory and experiments. J Siomech Eng, 102, 73
STRESS-RELAXATION RESPONSE
CARTILAGE IN COMPRESSION
viscoelustic stress-relaxation response
F,,,,,.t,r,"a, cartilage in a I D compression experidepi,ek'ci in Figure 3-10. In this case, a con,·,,,,,,rtression rate (line t,,-A-B or lower left figto the tissue until lin is reached;
b¢ypncl point B, the deformation Uo is maintained.
cartilage, the t.ypical stress response
this imposed deformation is sho\\'11 in the
flgurc (Holmes et aI., t 985; Mow ct aL,
1984). During the compression phase, the str
rises continuously until (To is reached, CO
sponding to lI o, while during the stress-relaxat
phase, the stress continuously decays along the cu
B-C-D-E until the equilibrium stress (U X ) is reach
The mechanisms responsible for the stress
and stress relaxation are depicted in the lower p
tion of Figure 3-10. As illustrated in the top d
grams, the stress rise in the compression phase is
sociated with fluid exudation, while stress relaxat
is associated with fluid redistribution within
porous solid matrix. During the compressive pha
Copious
fluid exudation
Fluid redistribution
(no exudation)
t 1I t
A
B
I
I
III
c
D
Equilibrium
deformation
E
I---"'L-o-a-d-in-g---ll jf------"S-tr-e-ss-'e-Ia-x-a-'i-o-n------
cw
B
E
w
~
a.
6
uo- - -
-
-
-
-
B C
D
E
;-~----....- ...- - I
o
Controlled ramp displacement curve imposed on a cartilage
specimen commencing at to (bottom left) and the stressresponse curve of the cartilage in this uniaxial confinedcompression experiment (bottom right). The sample is compressed to point B and maintained over time (points B to E).
The history of the stress and response shows a characteristic
stress that rises during the compressive phase (points to to B)
and then decreases during the relaxation phase (points B
D) until an equilibrium is reached (point E). Above these
curves, schematics illustrate interstitial fluid flow (represented by arrows) and solid matrix deformation during th
compressive process. Fluid e~udation gives rise to the pea
stress (point B), and fluid redistribution gives rise to the
stress-relaxation phenomena.
the hi!:!h stress is generated bv forced exudation of
the in~crstilial fluid and the c~mpaclion of the solid
matrix 11('<\1- the surface. Stress relaxation is in turn
causcd bv the relicf 0" rebound of the high COIllpaction r~gi(}n ncar the surface or the solid matrix.
This stress-relaxation process will cease when the
compressive stress developed within the solid ll1alrix
reaches the stress generated by the intrinsic COIllpressive modulus of the solid matrix corresponding
to U o (Holmcs ct 'II., 1985; Mow ct aL, 1980, 1984),
Analysis of this stress-relaxation process leads to lhe
conc"!usion that under physiological loading cOl1clilions. excessive stress levels arc difficult (() maintain
because stress relaxation will rapidly a(lenuat~ the
stress developed within the tissue; this must ncces~
sarily lead to the rapid spreading of the contact area
in tht:: joint during articulation (Atcshian el aI., 1995,
1998; tvlow & Atcshian, 1997),
Rcccntlv, much focus has been on the inholllogencitv o(I-IA with carLilagc depth (Schinagl ct aI.,
1996, '1997), Bascd on thi~ data, from an analysis
of the stress-relaxation experiment it was round
that an inhornogencous tissue would relax at a
faster rate than would the unirorlll tissue (\Vang &
1\110\V, 1998). iVloreovcr, the stress, strain, pressure,
and Ouid flow fields within the tissue were significanllv altered as well. Thus it seems thal the variation~' in biochemical and structural composition
in the layel"s of cartilage pl"ovide anothet- challenge
to understanding the environn1ent of chondroc.vtes in situ.
PERMEABILITY OF ARTICULAR CARTILAGE
Fluid-filled porous materials ma~' or may not be permcablc, Thc ratio of nuid volumc (VI) to thc total
volume (\IT) of the porous material is known as the
porosity (f3 :; ;: ViI VT); thus, porosity is a geOJ1lclric
concept. Articular cartilage is therefore a material
of high porosity (approximately 80%), If the pores
are int~rconnected, th.; porous material is permeable. PermeabilHv is a measure of the ease with
which fluid can f1(~W through a porous material. and
it is inversely proportional to the frictional drag ex~
ened by the fluid flowing through thc porous,
permeable material. Thus, permeability is a physical concept; il is a meaSure of the resistive force that
is required to cause the fluid to Ilow at a given speed
through lhe porous-permeable material. Thb fric~
tional resistive force is ~enerated bv rhe inleraction
of the intcrstitial fluid-and the p;'rc walls of the
porous-permeable material. The permeability coefficient k is related to the frictional drng coefficient K
by the relationship k ~ W/K (Lai & Mo\\', 1980), A
{j'cular cartilage has a \'(.'1:-' low !kTlllcabilily a
thus hi~h frictional l"I.... sistin= I"oret's arc g('n~rat
when n~,id is caused tn flo\\" through the poro
solid matrix.
In the previous sections on c;:\rtilage \·iscoclast
ity \\·e discw;st.'d the process of lluid !low throu
articular cartihH.!e induced bv solid matrix compre
sion and how ll~is process i;lfluences lI1L' viscoda
tic behavior or the tissuC'. This process ;:l1so provid
an indirect method to dctcrmine the permeability
the tissuc. In this section, we discuss the cxpc
mCJ1t~11 method uscd to direct I.\' measure the perm
ability codTicknt. Such an L'xperim(:nt is depict
in Fi~urL' 3-1 I.·\. Here, a specimen or the tissue
held fixed in ~\ chamber subjected to the action o
pressure gradicl1l: tilL' imposed upstrt.'al11 pl\:ssu
PI is greater than the dOwJ1slrt.:;:\Il1 pressure P2' T
thick~css 01" tlte specimen is denoted by han
the cross-sectional an..-a 01" pt.'nneatiol1 is ddin
b\· A. Darcy's law, L1sed to determine the perme~\b
it'\" k fron~ this simple t.'xpcrimental St.'ltlp, ~·id
k' = Qh/A(Pi-P2). where Q is the \'olumelric d
chargL' per unit time through the specimen who
area 01" permcm ion is A ([\llo\\' & R;:ltcl irr(" 1997). U
ing low pressures, approxiJ11atel~' 0_1 ivlPa. t
mClhod was f1rslused to determine the permeabil
of articular cartilage (Edwards, 1967; lVlnroud
1975). The value ~I" k obtained in this mann
ram!ccl from 1.1 x 10. 1 < m·l/N·s 10 7.6 ;:.< 10. 1 ' m·:/N
In ~~ddition, using a uniform straight tube mod
the ;:\\IL'rage "pore diameter" has been estimatt.:d a
11111 (Maroudas. 1979), Thus. the "pores" within
tinllar cartilage are of mokcular size.:',
The pL'nnt.'abilit~'or Hnicular cartibgc under co
pressive strain and at high physiological pressu
(3 MPa) was first obtained b~' j\Aansour and lV'1
(1976) and later analvzed by Lai and ,Vlo\\' (198
The high pressure and compressh"e strain conditio
examined in these studies more c1osd~' resem
those conditions found in dianhrodial joint loadin
In these l.'xperimcnts, k was measured as a functi
two variables: the pressure gradient across
specimen and the axial compressive strain applied
the sample. Tht.· results from these experiments
shown in Figure 3-11 B. Permeabilit~· decreased exp
nentially ~lS a function of both increasing compr
sive slr~lin and increasing applied Iluid pressure,
was hlter shown, however, that. the dependence o
Ion the applied fluid pressurt.: derives from co
paction of the solid matdx that. in lurn, results fr
the frictional drag ("used by rhe permeating nu
(Lai & MO\\', 1980), From the point of view of p
or
Fluid
pressure
P
Rigid
porous
blocks
,,
{
Fluid
pressure
P2
'-
Fluid flow
"~
"e
I
i
~
C><I;:,
.. ,;.;.;+:.:.:. :.:.:-: .;.:.:....
'-
J
-
I
, , ,
,
- I, 1Fluid1flow
I I
• • ! !
~
I
-'--
0
h
L
~
x
•.
:c
Articular
'"
§'"
cartilage
-UIlIU
..I
. ,1
~
A. Experimental configuration used in measuring the perme·
ability of articular cartilage, involving the application of a
pressure gradient {P,-Pl)/h across a sample of the tissue (h =::
tissue thickness). Became the fluid pressure (PI) above the
sample is greater than that beneath it (PJ, fluid will flow
through the tissue. The permeability coefficient k in this experiment is given by the expression Qh/A(PI-P/), where Q is
the volumetric discharge per unit time and A is the area of
permeation. Adapted from Torzilli. PA, & MoJ.v, VC (976). On
rhe fundamental fluid cranspon mechafllsms rhrough normaJ and
a.'"
Applied
pressure difference
16
14
(P,-P,)
r
10
... 0.342 MPa
r
00.689 MPa
~
8
1.034 MPa
01.723 MPa
6
4
2
r
o[
0
B
.0.069 MPa
.0.172 MPa
12
8
16
24
32
40
Applied Compressive Slrain {%}
parhologic caail,1ge (Juring function. I. The formulation. J Bio-
meeh. 9(8), 5:11-552, a, Experimental curves for articular cartilage permeability show its strong dependence on compressive strain and applied pressure. Measurements were taken
at applied pressure differential (p\-p)) and applied strains.
The permeability decreased in an exponential manner as a
function of both increasing applied compressive strain and
increasing applied pressure. Adapted from L,1i, 1,1I/,'\11.. & MoV'-/,
II C. (/980). Drag-induced compre5sion of ,1rtiwlar (,milage during a pe-rmeaiion experimenr. J Btorheology, 17. J 11 .
~---------------------------------
';S,n~IC[lIre, compaction or the solid matdx
9j~ porosity' and hence the average "pore
decreases
diameter"
" \\Iithin the solid matrix; lhus, solid matrix compaction increases fTictional resistance (MOWCl aI., 1984).
... >",:1'he nonlinear permeability of articular cartilage
'X de.monstrated in Figure 3-11 B suggests that the lis<~/!7:9~ has a mechanical feedback sYSlem that may
.J;s,e~'ve important purposes under physiological con:'o'ditions. When subjecled 10 high lo"els through the
\mechanism of incrcnsed frictional drag against in:)~rslitial l1uiel 110w. the tissue will "ppear sLirfer "nel
:';L'h will be more difficult to cause fluid exudation. Re·
(cent analyses of articular cartilage compressive
/,',/ ':,s,trcss-relaxation behavior have validated this con;C:"'c'ept and its importance in the capacity of the interg, stili,,1 Ouid La support load (Ateshian eL al.. 1998:
;'::;',":,~oltz & Ateshian, 1998). Moreover, this mechanism
;< /al~o is important in joint lubrication,
/9BEHAVIOR OF ARTICULAR CARTILAGE
§;&,UNDER UNIAXIAL TENSION
~~?''.::~:;rhe mechanical behavior of articular canilage in
3~~;'{enSion is highly complex. In tension, the liss~le is
,:~.~;
A
strongly anisotropic (being stiffer and stronger for
;'::;;--"
'.C:;
'i";
".'.,
specimens h"rvested in the direction parallel to the
split line pattern than those harvested perpendicular to the split line pattern) and strongly inhomogeneous (for mature anirnals. being stiffer and
stronger fOl' specimens harvested from the superficial regions (han those harvested deeper in the tissue) (Kempson. 1979: Roth & Mow, 1980), Intercstingly, articular cartilage rrom immature bovine knee
joints does not exhibit these layered inhomogeneous variations; however, the superficial zones o
both ITlature and immature bovine cartilage appear
to h"ve the ,,,me tcnsile stiffness (ROLh & Mow
1980). These anisotropic and inhomogeneous characteristics in mature joints arc believed to be caused
by the vaI)'ing collagen and PG stnlctural organization of the joinl surface and the layering structural
arrangements found within the tissue_ Thus, the collagerH'ich superficial zone appears to provide the
joint cartilage with a tough wear-resistant protective
skin (Sellon et "I.. 1993) (Fig. 3-3;\),
Articular cartilage also exhibits viscoelastic behavior in tension (\-Voo ct "I.. 1987). This viscoelastic behavior is allributable to both the internal friction associated with polymeric motion and the flow of the
interstitial fluid. To examine the intrinsic mechanica
-i·
response of the collagen-PC solid matrix in tension,
it is necessary to negate the biphasic fluid now effects. To do this, one mllst perform slow, low strainrate experiments (/\kizuki et aI., 1986; Roth & Mow,
1980; Woo et aI., 1987) or perform an incremental
strain experiment in which stress relaxation is al-
;.:
;;.
lowed to progress IowaI'd equilibration at each incremenl of strain (Akiwki et aI., 1986). Tvpically, in a
low strain-rate (or near-equilibrium tensile) experiment. a displacement rate of 0.5 em/minute is L1sed
and the specimens usually arc pulled to failure. Unfortunately. lIsing these procedures to negate the ef~
feet of interstitial Ollid now also negates the Illanires~
tation of the intrinsic viscoelastic behavior of the
solid matrix. Thus, only the equilibrium intrinsic mechanical propenies of the solid matrix ma~r be determined from these tensile tests. The intrinsic viscoelastic properties of the solid matrix Illust be
determined from a pure shear study.
The "equilibrium" stress-strain curve ror a specimen of articular cartilage tested under a constant
low strain-rate condition is shown in Figure 3-IL
Like other f,brous biological tissues (Iendons and
ligaments), articular cartilage tends to stiffen with
increasing strain when the strain becomes large.
Thus, over the entire range of strain (up to 60 0m) in
tension, articular cartilage cannot be described b~: a
single Young's modulus. Rather, a tangent modulus,
defined by the tangent to the stress-strain CUI'\'C.
must be used to descl'ibe the tensile stiffness of the
tissue_ This fundamental result has given rise to the
wide range of Young's modulus. 3 to 100 ,\lIPa, reponed for anicular cartilage in tension (Akizuki et
aI., 1986; Kempson, 1979; Roth & Mow, 1980; Woo
et aI., 1987). At physiological strain levels, however,
less than 15% (Armslrong et aI., 1979) of the linear
Young's modulus of anicular cartilage ranges between 5 and 10 MPa (Akiwki et aI., 1986).
Morphologically, the cause for the shape of the
tensile stress-strain curve for large strains is depicted in the diagrams on the right or Figure 3-12.
The initial toe region is caused by collagen fiber
pull-out and realignn'lCnt during the initial pan ion
of the tensile experimel1l, and the final linear region
is caused by the stretching of the straightcnedaligned collagen fibers. Failure occurs when all the
collagen fibers contained within the specimen are
ruptured. Figure 3- I 3/\ depicts an ullslretched articular cartilage specimen, while Figure 3-138 depicts
a stretched specimen. Figure 3-14, A & B shows'
scanning eleclron micrographs or carlilage blocks
under 0 and 30% sU'elch (right) and the COlTesponding histograms of collagen fiber orientation
+i TenSIle modulus. E
I
I
'=~
Ci!t
-
~~~
Failure
c
.; 1
.!----
~
2
iii
i
~--\
~~~
Linear regio
I
I
.-------
L
,=------------~
~
~~~
Toe region
~_S-'-'''-;n-.-'_(_~--U-L-O-)------
Typical tensile stress·strain curve for articuliH cilrtilage
dril\Nings on the right of the ClJrve shm.", the configu
of the collagen fibrils al vcUiOllS stages loading. In th
region. collagen fibril pull·out occurs c15 the fibrils ali
themselves in the direction of the tensile load. In the
eM regiol1, the aligned collagen fibers are stretched
failure occurs
.--i
------------
delermined from tilL' ~cnnnil1g dcctron microg
pkltln..: s (h.'h). Ch.'arl.\· it can he seen lhal ll1l.'
gcn lletwork within c~lrlil~\gL.' l"L'sponds to le
sll\.. s~ and strain (\Vada 1..'\:. Akizuki, 19~7l.
If the mok'ctll~lr st!1H.:(urc or L'(Jllag('~Il, Ilk' orga
lion of thL' collagen fibers within the collagclloll
\\"01''', or tilL' collngl'1l IiIK'!' lToss-lillking is al
(such ~lS tin\! OcclIlTing in mild fibrillation or OA
tensile prolKTlies of the I1L'I\\'ork \\'ill ch~lIlge. Sch
l't al. (1990) ha\'c shown a definitive' t\:.-btionsh
1\\"C'I.'11 collag.. . n hydroxypyridinillm cn)ss-linkin
lensik' stilflk.'SS ~II)(I stn.. ,lgtll of nOI'mal h(wirH.~
hlg(·. Aki/.uki et al. (1086) showcd lhat pr()grc
dcgr~\dnlion or lUI man kllL"C joint cartilag:(.', from
Of\. ~'iclds n progn..:'ssin· (k'kriorati
propl".'1'liL'S IJI" Ihe collagcn-PG
matrix. Similar rl.'sults 11<.1\'(' bccll obsen'\..-d n.. .c\..'
'lI1ill1"lll1odcls of O!\ (Guil", l't ,Ii.. 1004; SCt!(Hl
It)t).t). l()g\..,tllcr, lhese obSCI·V~llioJls sllpporl tll'..'
that disruptioll of tilL' colbgcn IlL'lwork is ~\ kL'~'
ill thL' initial CH'IHS iL'adillg 10 11K' dC\'L'lopmcllt o
Also. loosening of til\.' collagen network is gCI1\..'ral
librillation
10
Ilk' illl rinsic Icnsik'
licved to h(' responsible for tlte incrcased swc
hCllci..' W~Il\..T COI1H.'nl. or ostL'oartlll'ilic car
(Mankin &: Thrasher, 1973: iVlaroudas. 1979). \,Vc
;:d]"L'ad~' discussed htl\\" iJ1cl"L'asl'd walcr cOlllClll
Unloaded
..AIf--+ Collagen
+--f-water
Proleoglycan
i
'
E.R., Lal: V:l.M.• & Mow V.C (1984). A continuum rheory and a
cartirage. J Bio
experiment for the ion-induced swelling behavior
mech Eng, 106(2), 15/-158
to decreased corripresstve stiffness and increased perrneability of arti~ular cartilage.
BEHAVIOR OF ARTICULAR CARTILAGE
IN PURE SHEAR
In tension and cOIllpression. only the equilibrium
intrinsic properties of the collagen-PG solid malri~
can be determined. This is because a volumetric
change always occurs within a material when it is
subjected to uniaxial tension or compression. T
volumetric change causes interstitia) nuid now a
induces biphasic viscoelastic effects within the
sue. H, however, articular cartilage is tested in pu
shear under infinitesinlal strain conditions.
pressure gradients or volumetric changes will
'produced within the material; hence, no interstit
fluid flow will occur (Hayes & Bodine, 1978; Z
et aJ. 1993) (Fig. 3-15). Thus, a stead v dynam
pure shear expct'iment C£ln be used [0 assess
Tension 0%
•
501
40
1
C
n=203
x = 52.0' " 23.0'
;
~ 30 1
20-i
£
1
~ lmmJlLill.IILillilllllLffi,.
o
45
90
Degrees
A
Tension 30%
•
50 i
n = 145
J:::~!
'8.9~
'km= ::
lOi0-'
o
I
17.6':>
In,m_r!L~
__,.
45
90
Degrees
B
Direction of Load
Collagen fibril alignment is dearly demonstrated by the scanning electron micrographs
(X10.000) (right) of cartilage blocks under 0% stretch (A) and 30% stretch (8). The histograms (left), calculated from the micrographs. represent the percent of collagen fibers
oriented in the direction of the applied tension. At 0% stretch the fibers have a random
orientation; however. at 30% they are aligned in the direction of the applied tension.
Reprinred with permission (rom Wada, T.. & Akizuki, S. (1987). An ulcrasuucllIral sllIdy of solid matrix in articular cartilage under uniaxial tensile stress. J Jpn Orthop Assn. 61.
intrinsic viscoelastic propcrLics or the collagen-PG
solid matrix.
In a steady dynamic shear experiment. the viscoelastic properties of the collagcn-PG solid matrix
are detenl1ined by subjecting a thin circular wafer
of tissue to a steady sinusoidal torsional shear,
shown in Figure 3-16. 1n an experiment of this type,
the tissue specimen is held by a precise amount of
compression between two rough porous platens.
The lower platen is attached to a sensitive torque
transducer and the upper platen is attached to a
precision mechanical spectrometer with a senrocontrolled de mot01: A sinusoidal excitation signal
may be provided by the motor in a frequency of excitation range of 0.01 to 20 hertz (Hz). For shear
strain magnitudes ranging from 0.2 to 2.0°10, the
viscoelastic propcnies arc equival~nl1y defined
the elastic storage modulus G'; the viscous
modulus G" of the collagen-PG solid matrix ma
determined as a runction of frequency (Fling, 1
Zhu et aI., 1993).
Sometimes it is morc convenient to determ
the magnitude of the dynamic shear modulus
given by:
IG*I' = (G')' + (G")'
and the phase shifl angle given by:
l)
= tan'
(G"/G')
The magnitude or lh~ d~'namic shear modul
a measure of the tala I resistance offered bv the
coelastic material. The value of 0, lhe angl~ betw
[
I
I
Unloaded
Pure shear
I
Collagen
+----fWater
Proteoglycan
B
Schematic depiction of unloaded cartilage (A), and cartilage subjected to pure shear (8). When cartilage is tested in pure shear
under infinitesimal strain conditions. no volumetric changes or pressure gradients are produced; hence, no interstitial fluid flow
occurs. This figure also demonstrates the functional rote of collagen fibrils in resisting shear deformation.
','the steady applied sinusoidal strain and the steady
~ipusoidallorqllcresponse,
-.0,"
is a measure of the total
':";Edctional energy dissipation within the ll1aterial.
EoI' a pure elastic material with no internal fric,:'ti'onal dissipation, the phase shift angle 8 is zero; for
:iijJurc viscous fluid, the phase shih angle 8 is 90t '.
)/;·The magnitude of the dynamic shear modulus for
'i':t1()rmal bovine articular cartilage has been measw\:d to range from I to 3 ~llPa, while the phase
'r shift angle has been measured to range from 9 to
20' (Hayes & Bodine, 1978; Zhu Cl aL. 1993), Thc inirinsk transient shear stress-relaxation behavior of
::':tI1t~ collaucn-PG solid matrix along with the steadv
'\Iynamic . . .s hear properties also ha; been llleasUl'e~1
;',:. (Zhu el aI., 1986), Wilh bOlh lhc sleady dynamic and
'::'thc. transient results, the laller investi'gators showed
. ~at the quasilinear viscoelasticity th~ory proposed
'?;:'.py Fung (198l) for biological materials provides an
<"accurale dcscription of the flow-independent visi; coelaSlic behavior of lhe colla!!cn-PG solid malrix,
'.:':/,}<'jgure 3-17 depicts a comparis--on of the theoretical
'prediction of the transient stress-relaxation phe;[lomenon in shear with the results from Fung's t 981
'quasilinear viscoelasticity theory.
/:.. From lhese shear sluciies. it is possible to obrain
SOme insight as to how the collagen-PG soHd matrix
functions. First, we note that measurements of PG
>:. SOlutions at concentrations similar to those found in
':>a. nicular cartila(Tc
eo in Silu .'vidd a ma~miLudc of shear
0:;"
modulus 10 bc of thc order of 10 Pa and phase shifl
angle ranging up to 70" (Mow et aI... 1989b; Zhu ct
aI., 1991, 1996), Therefore, it appcars that the magnitude of the shear modulus of concentrated PG so-
SIN (f
n
I
'd
I
..
111
I
I
~I
Time
Steady sinusoidal torsional shear imposed on a specimen in
pure shear. The fluctuating strain in the form of a sine
wave with a strain amplitude filand frequency f.
Iht: ullla~I..·1l 111..'1\\1 Irk. 1')I,.·rJllib till..' PC ~I.
to rl..':--.i:--t c{JlI1l)rl..·.... ~ion (
192-l: Marolld.. I:--'. 1~79: ,'vl0\\·I..\: Rall.:lilll', 1997
cOlllll for :--.tH.'h Fix('d Char~l..· DI..'I1:--.ily (FeD, 1
cartila1!l'. a lriplla~i(.· Ilk,dlancl-l..·kdnl\,:llelTlica
1..·!L·I...·ln,I.\·k· Ihcor,\' \\":I~ d ...·\·t.. !opl..·d [kit rJl(ldd~
"l~ a mi'\tlIrl...·
Ihn..·\,· misdhk' pklsl..·s: a cklr~l
pkhl..· rl..'j)rL·SI..·llIill~ IhL' \.:(jIl"I~I..·n.PC IlL'I\Uli
phasL' rl..'l.,rL·~I..'llling th\,o inll..'rstiti .. d \\·<llI..'r, <lll
ph"lse cOlllpri~illg tlk' llltlnO\·i.lk'llt L'alillil ' a ion CI a~ \\"1..,11 as Dllk'l' 1l111llh·.. dl..·1l1 Slk'l..:iL's
C"I' (ClI L'I <II., llJlJ:{: l.'li 1..'1 al..llJ01 ). I Il I hi~ I h
IOlal slrl..·s~ i:- ~in'll h.\· IIIL' slim
1\\'0 1I..'rllls
CT,,,h.1 .~. Ulh"'i. \\"hel·I..' (T,,·"·I .llId (T,:',,·I "Irl' Ihl' ~
ill
cc.,II"I~l:1I Ill't\\"ork
1.0
11
i5'
C
·uc0
'=
0.. 0004
12"" 36.2
c = 0.13
0.8
~
(lr
U.
c
.Q
1ii
x
~
!
a;
c: 0.'
j
!
j
!
!
'0
~
U
~
.,
'0
or
a: 0.2
0.0
0
10
20
30
Time (sec)
Typical stress-relaxation curve after a step change in shear
strain. expressed in terms of the mean of ten cycles of
stress relaxation normalized by the initial stress. The solid
line represents the theoretical prediction of the quasi lin·
ear viscoelasticity theory, Adapted from lilu, WB .. Lai. WA1.,
& Mav/, Vc. (/986). Intrinsic quasi.. /inear viscoelasric beflMior
of the extracellular matrix of cartilage. Trans Orthop Res SoC,
11,407.
);
IUlion is one hundred thousand times less and the
phase angle is six to seven times grcmcr than lhat of
anicular cartilage solid matrix. This suggests lhat
PGs do nOl function in situ to provide shear stiffness
for articular carlilage. The shear stiffness of arLicLIlar cmotilage Illust lherefore derive (Tom its collagen
content, or from the collagen-PG interaction (Mow
& Ralcliffe. 1997). From Ihis interpretation, an increase in collagen, which is a much more clastic element than PC and the predominant load-carrying element of the tissue in shear. would decrease the
frictional dissipation and hence the observed phase
angle.
SWELLING BEHAVIOR OF ARTICULAR
CARTILAGE
The Donnan osmotic swelling pressure, associated
with the densely packed fixed anionic groups (SO,.
and COO·) on the GAG chains as well as the bulk
compressive stillness of the PG aggregates entangled
tri.'\ sln:ss <Inti inli:rsliliailluid pr\'·~:--.un... r",,:-I..·quilibl·illill. (T'l,,,,, is gi\·I..·!l h~' I hI..' [)OIlIl ..Hl
prcssur..... IT hcl' disCllS~i(J1l Ik·lu\\·l. J)erin'd
till..' fulltl ..t1n.... nl .. t1 laws (II' IllI..'C!J"lIlics <.llld
d.\"ll~tlllics rather Ill"lll through IIh..· ad 11m: c
tion of L'xisting :--pi..·i.:i ..l1i·I.I..'d thl·{Iril...·:-- k.g ..
Grod/.insk.\", [YS/a,hL this 11'ipkl.... ic Ilk'or,\"
~\ ",\..-'1
I!lLTln(ldYIl~IJllicalh' Ik·rllli ...... ihk C
li\'\..· law.. . to dl''''lTibl..' lit\..-' till'lL'-dL'[k'lllk'llI
l·lll..'lllic~d, Ill\..-'Cll~lnil·;d, :IIHI L'k'l'lriL"al pI'OpL
cll<:II'gL'd-!lH!ralL'd so!'t tiSSlll
;\'11. 1I'L'O\·\..-'1', Ill\..
."ic 11lulti"ek'ctrol\'!L' tlwol',\' h<l h l 'L'1'l shown
tirl·l.\' l'on ... istl'11{ \\'ith thl' ",pl'ci<tlih·d clas
nWlic PI'I..'SSUl'L' thL'(lly fut' ch'll'~l'(.1 POIYnll.'
ti011S, plh.'llOllll·llologic;,,1 {1'~lnSpol't II1L'ori
tilL' hiphask theor.\· (DOIlIl .. IIl, 192-l: I(;'ltch
Curran. 1973: .\'Io\\' 1..'1 .. 11 .. 19~O: On... a~L'r. 193
\\'hkh In\\''''' bL'I...·1l frL'lI11L'ntly llsL·d 10 slud.\·
f.. lcl..'ls (II' aniCllbl' clrlilagL·.
Th...' Iriphasic IhL'or~' has bl...·I...·!1 used sliccl...
describe lll .. tny clf till..' 1l1L'l'll"111(I-l'kcII'Oclll..
ha\·ior:-. or articubr cartilage. Tl1l..'s...· include
diclion
frl..·e S\\·L·lIing lIIHk'r c1l1..'lllicalload
L'a!' dCIK·IH.k·IlCL· or h.nlr"llllic pl·rnll..·i.lhilil.\· \\"
llo!llirH.·;11' depcn(k'llCC
stl'L'allling pOIL'llt
FeD: cLll'ling or Glnil ..lg,l· l"l.\·LT:-: pn.'-qrl'ss:
..Inc! Ilcgatin' osmotic flows: s\\'clling ..Ind l
responses or cells 10 11SIlH>lic SIH)(.:k IO<lding;
inlllll'lln:.' or inhomog,('IlL'OUS fi.'\l·d dl<:lr~c
(CUd al.. 1993. 1997. 1<.)%; Lai ,'1 "i.. 1<,)<,)1
al.. J 998; Sellon L·t al... 1Y9t': SUIl ('I aI., 199
\'idillg Illorl..' \·I..'rsalilit.\·. Ih...• triphasic Ih...·ol:·
gcneralizcd to indlldl' lllulti-dcCII'OI.\·ICS in I
(Gil L'I al.. 1Y9~).
From analysis llsin~ !l1l' triphasic theor.
cOllles clear thai thl' swelling hdw\'lor or lh
(:an hI..' I'L'sponsibiL- for ..I sigllific<IlH fractio
compn,:ssi\'c IOi.ld-bL.'~lriJlg Ci.lp~\i.:it~'
artic
..-\1
(lr
ur
or
or
or
tilagc at equilibriulll (Mow & Ratcliffe, 1997). For
example, the triphasic theory' predicts for confined·
compression at equilibrium that the total stress
((Jl"I'd) acting on the cartilage specimen is the SlIlll
of the stress in the solid matrix (U"'lid) and the Don·
nan osmotic pressure (Ulh'id = n). The Donnan osmotic pressure is the s\\.'clling pressure caused by.'
the ions in association \vith the FeD and represents
the ph:vsicochcmical motive force for cartilage
~\Vcl1ing (Fig. 3-18). From the classical theOl}' for
osmotic pressure, the Donnan osmotic pressure
caused by the excess of ion pnrticlcs inside the tissue is given by::
IT
= RT[<[,(2c+c')-2<!,*c*j +
p=
where c is the interstitial ion concentration, c'" is the
external ion concentration, c" is the FeD, R is the
universal gas constant, T is the absolute temperature, (t) and 4)'" are osmotic coefficients. and peo is
the osmotic pressure caused by the concentration of
PG particles in the tissue, usually assumed to be
negligible (La! et aI., 1991). For a lightly loaded tisSll~, ~he s\vclling pressure ma:v co~~tribule significantly to the load support. But for highly loaded tissues, such as those found under physiological
conditions and certainly.' for dynamically loaded tissues, the interstitial fluid pressurization (Jtlllid)
\vould dominate; the contribution of this swelling
pressure to load support would be less than 5(jf;
(Soltz & Ateshian, 1998).
As with the biphasic theory, the triphasic n1echano~
electrochemical theory can be used to elucidate
potential mechanosignal transduction mechanisms in cartilage. For example. because of their
potential effects on chondrocyte function, it is important to describe and predict electrokinetic phenomena such as streaming potentials and stream~
ing currents (Gu et aI., 1993, 1998; Katchalsky &
Curran, 1975; Kim et al" 1994) that arise from ion
movement caused by the convection or interstitial
fluid flow past the FCD of the solid matrix. As a
second example, the pressure produced in the interstitial fluid by polyethylene glycol-induced osmotic loading of cartilage explants (Schneiderman
et aI., 1986) was recently shown to be theoretically
nonequivalent to the pressure produced in any
other common Iv used mechanicallv loaded explant experime~t or by hydrostatic -loading (Lai
et aI., ] 998). In light of this finding, earlier interpretations of biological data from studies making
such an assumption of equivalency should be reVisited.
0.4
Swelling pressure
rr
0.3
0.2 -
0.1
o
0.15
0.5
1.0
Bathing Solution Concentration c' (M)
Swelling pressure of articular cartilage versus bathing s
tion concentration (c*). At equilibrium, the interstitial f
pressure is equal to the swelling pressure, which is defin
by the tissue Donnan osmotic pressure (IT).
Lubrication or Articular
Cartilage
As alread.y discussed, synovial joints are subjec
to an enormous range of loading conditions,
under normal circumstances the cartilage surf
sustains little weac The minimal wear of norm
cartilage associated with such varied loads indica
that sophisticated lubrication processes are at w
within the joint and within and on the surface of
tissue. These processes have been attributed to a
bricating fluid-film forming between the articu
cartilage surface and to an adsorbed boundary
bricant on the surface during motion and load
The variety or joint demands also suggests tha
number of mechanisms are responsible for
arthrodial joint lubrication, To understand diarth
dial joint lubrication, one should use basic e
neering lubrication concepts,
From an engineering perspective, there are
fundamental types of lubrication. One is bound
lubrication, which involves a single monolayer o
bricant molecules adsorbed on each bearing surfa
Subchondral bone
d;J <:>\.kV
(Q:>
~ Load
Articular cartilage
Synovial fluid gap
Articular cartilage
A
Subchondral bone
Hydrodynamic
d;J <:>\.kV
yn
n
(Q:>
d;J <:>\.kV
n
Load and
Motion
Load and
yMo/ion
~
Load and
yMotion
• p
+f-o
•
B
Squeeze-Film
c
Boosted
A, In hydrodynamic lubrication, viscous fluid is dragged
into a convergent channel. causing a pressure field to be
generated in the lubricant. Fluid viscosity. gap geometry,
and relative sliding speed determine the load-bearing capacity. B. As the bearing surfaces are squeezed together,
the viscous fluid is forced from the gap into the transverse
direction. This squeeze action generates a hydrodynamic
pressure in the fluid for load support. The load-bearing (a-
The other is fluid-film lubrication, in which a thin
fluid-filn1 provides greater sllrrace-to~sllrraceseparation (Bowden & Tabor; 1967). Both lubrication types
appear to occur in articular canilage under vaJying
circulllstances. Intact synovial joints have an extremely low cocrf!cicnl or friclion, approximately
o
Weeping
pacity depends on the size of the surfaces. velocity of a
proach, and fluid viscosity. C. The direction of fluid flow
under squeeze-film lubrication in the boosted mode for
joint lubrication. D. Depicts the Weeping lubrication hypothesis for the uniform exudation of interstitial fluid f
the cartilage. The driving mechanism is a self-pressuriza
of the interstitial fluid when the tissue is compressed.
0.02 (Dowson. 196611967; Linn. 1968; McCutch
1962; Mow & Atcshian. 1997). Boundary-lubrica
sllIfaces typically have a coefficient of friction one
two orders of magnitude. higher Lhan surfaces lu
cated by a fluid-111m, suggesting LhaL synovial joi
are lubricated. al Icast in part. by the fluid-f
?,\/,.. :.': :t'ni~~Aanism. It is quite possible that synovial joints
;;{',.~. ::.'\i~.{tihe:;mechanism that will most effectively provide
;ft\:.
fU$rica~ion at a given loading condition. Unresolved,
.;~:;' ~23:Jnc}Y:fllFjS the manner by which synovial joints gcn-
';%.;0;;:~~;>tethe fluid lubricant film.
;/:;
;:';..c;.
;,~Z
""''',.'.
~~r
'
}-iN,., Fti.JIP-fILM LUBRICATION
;;;.;;iir$Jdi~i!ri]m
lubrication utilizes a thin film of lubl'i"H:~'~'.> c~nt that causes a bearing surface separation. The
'6.·Jld.~a on the bearing is then supported by the pres";j~\.tt·~that is developed in this fluid-film. The nuid.'";:' ,..~';dihn';thickness associated with engineering bearings
C/. 'isitsually less than 20 f.l.m. Fluid-film lubrication re' ... q"jres a minimum nuid-fllm thickness (as predicted
,i':i>Y/a"speciflc lubrication theOIy) to exceed three
~f~;tioles' the combined statistical surface roughness of
cahUage (e.g., 4 to 25 f.l.m; Clarke, 1971; Walker et
,al:;)970). If fluid-film lubrication is unachievable
, ',;':,b~C~lllse of heavy.. and prolonged loading, incongru'e'nL'gap geomclI)', slow reciprocating-grinding 1110tion, or low synovial fluid viscosity, boundary lubrication must exiSl (Mo\\' & Ateshian, 1997).
The two classical modes of fluid-film lub"ication
" defined in engineering are hydrodynamic and
squeeze-film lubrication (Fig. 3-19, A & B). These
modes apply to rigid bearings composed of relatively undeFormable material such as stainless steel.
l-lydrodynamic lubrication occurs when nonparallel
rigid bearing surfaces lubricated by a nuid-film
move tangentially with respect to each other (i.e.,
slide on each other), forming a converging wedge of
Ouie!. A lifting pressure is generated in Ihis wedge by
the fluid viscosity as the bearing motion drags the
fluid into the gap betwcen lhe surfaccs, as shown in
Figure 3~ 19A. In contrast, squeezc-f-Ilm lubrication
Occurs when the bearing surfaces 1110\'e perpendiculad.''''- toward each other. A pressure is generated in
the fluid-film as a result of the viscous resistance of
the fluid that acts to impede its escape from the gap
(Fig. 3-19B). The squeeze-film mechanism is suffJcientto cany high loads for short durations. Eventually, however, the fluid-film becomes so thin that
contact between the asperities (peaks) on the two
bearing surfaces occurs.
Calculations of the relative thickness of the fluidfilm layer and the surface roughness are valuable in
establishing when hy'drodynarnic lubrication may
exist. In hydrodynamic and squeeze-film lubrication, the thickness and extent of the nuid-fllm, as
well as its load-carrying capacity, are characteristics
independent of the (rigid) bearing surface material
properties. Thcse lubrication characteristics are in-
stead determined by the lubricant's properties, such
as its rheological properties, viscosity and elasticity
the film geometry, the shape of the gap between th
two bearing surfaces, and the speed of the relativ
surFace motion.
Cartilage is unlike any man-made material with
respect to its near Frictionless properties. Classica
theories developed to explain lubrication of rigid
and impermeable bearings (e.g., steel) cannot fully
explain Ihe mechanisms responsible for lubrica
tion of the natural diarthrodial joint. A variation
of the hydrodynamic and squeeze-film modes o
fluid-film lubrication. for example, occurs \vhen
the bearing material is not rigid but instead rela
tively soft, slich as with the articular cartilage cov
ering the joint surface. This type of lubrication
termed elastohydrodynarnic, operates when th
relatively soft bearing surfaccs undergo either
sliding (hydrodynamic) or squceze-film action and
the pressure generated in the fluid-film substan
tiallv deforms the surfaces (Fig. 3-19, A & B)
These deformations tcnd to increase the surfac
area and congnlcl1cy. thus beneficially alterin
film geomCli)'. By increasing the bearing contac
area, the lubricant is less able to escape from be
tween the bearing surfaces. [t longer-lasting lubri
cant film is generated. and the stress of articula
tion is lower and 1110re sustainable. Elastohydro
dynamic lubrication enables bearings to greatl
increase their load-carrying capacity (Dowson
196611967,1990).
Note that several studies have shown tha
hyaluronidase treatment of synovial fluid, which de
creases its viscosity (to that or saline) b~; causing de
polymerization of HA, has lillie effect on lubricatio
(Linn, 1968; Linn & Radin, 1968). Because nuid-fllm
lubrication is highly dependent on lubricant viscos
ity, thcse results strongly suggest that an alternativ
1110de of lubrication is the primary mechanism re
sponsible for the low frictional coefficient of joints.
BOUNDARY LUBRICATION
During diarthrodial joint function. relative motio
of the articulaling surfaces occurs. In boundary lu
brication. lhe surfaces are protected by an ad
sorbed layer of boundary lubricant, which prevent
direct. surface-to-surface contact and eliminate
most of the surface wear. Boundarv lubrication i
.' essentially independent of the phy~ical propertie
of either the lubricant (e.g .. its viscosit~,) or th
bearing material (c.g., its stiffness). instead de
pending almost entirely on the chemical propertie
of the lubricant (Dowson, 1966/67). In synovial
joints. a specific glycoprotein. "Iubricin," appears to
be the synovial fluid constituent responsible for
boundary lubrication (Swann ct aI., 1979, 1985).
Lubricin (25 X 10·: In\\') is adsorbed as a macromolecular monolayer to each articulating surface
(Fig. 3·20). These two layers. ranging in combined
thickness frol11 I La 100 11m, arc able to carry loads
and appear to be effective in reducing friction
(Swann et aI., 1979). More recently, Hills (1989)
suggested that the boundary lubricant found in
synovial fluid was more likely to be a phospholipid
named dipalmiloyl phosphatidylcholine. Allhough
experiments demonstrate that a boundary lubricant can aCCOlint for a reduction of the friction coefficient by a factor of tbreefold to sixfold (Swann
et aI., 1985; Williams et aI., 1993), this reduction
is quilc modest corn pared wiLh the much greaLM
cr range (e.g., up to 60 fold) reported earlier
(McCutchen, 1962). Even so, these results do suggest that boundal)1 lubrication exists as a complementary mode of lubrication.
M
-_._--
MIXED LUBRICATION
There are two joinl lubrication scenarios thal can be
considered a combination of fluid·fiIm and boundary
lubrication or simpl~' mixed lubricalion (Dowson,
1966). The first case refers to the temporal coe\istence of fluid-film and boundary lubrication at spatially distinct locations, whereas the second case,
termed "boosled lubrication," is chanlclerized b~" a
Scanning electron miuograph of the surface of huma
articular cartilage from a normal young adult showing
the typical irregularities characteristic of this tissue
(": 3,000). ;.ldap;ed from Armsrrong. C G.. 8 t.J1o'.·.~ V C
fl980r Fncrlon. fu!mCM!ol) dfld \~/ear of Synovkll Jomrs. In'
O'."If.'fJ. )
Goodfe!lo~·'1.
<1.'1(1 P B(Jl!ouqh {Eds.i. Scientl1,( Fou
Traumatology fpp 223-232j Lo
don Wilham Neinermcl!lft
"ons of
Or;hopa~dlcs cwd
.----------------Articular surface
~n~1\\l~lf2t±-
Lubricating glycoprotein
Articular surface
Boundary lubrication of articular cartilage. The load is
carried by a monolayer of the lubricating glycoprotein
(lGP). which is adsorbed onto the articular surfaces. The
monolayer effectively serves to reduce friction and helps
to prevent cartilaginous wear. Adapred from Armsrrong.
C G., & Mo'."/, v: C (1980). Friction, rubrication and ~'/ear of
synovial joints. In: R. Owen, J. Goodfellow, and P Bullough
(Eds.). Scientific Foundations of Onhopaedks and Traumatology (pp. 223-232). Londofl: Wilfiilm Heinermann.
•
shirl or 1111id-film lo bOl1ndar~-" luhrication with l
0\'(.'1' the same loc'-llion (\Valkcr I.'l aI., \070).
The articular cartilagl...' :-;urf<'KL~. like all surface
not pL'rfCCtl.v smooth; as(k'ritics project out from
surface (Clarke, 1971; Gardner oS: McGillivray. 1
Rc(IIl..~r& Zimn~', 1970) (Fi~s" 3-3/3 and 3·21). In s
ovialjoints. situations may occur in which lhe ll
film thickness is or the san'll.' order as the 111('..\11
ticular surface aspt..'rity nV..tlkcr ('t .. II", 1970), Du
such insl . Hlct..'S, b()undar~' lubrication betweell
asperities may COllle into pla~·. If this occurs
mixed mode of lubric..ltinn is npcr"lting, with
joint surface load sustailk'd b~' both the lIuiclpressure in "l1"eas
nOIlCOlllact .. lIlel b~' the bCH
..Iry lubricant lubricin in the .. m:.'as of i.lspcrit.\' c
tact (showll in Figure 3-22)" In this mode of mi
luhl"icntion. it is probable lit..\( Illost 01" the fric
or
(which is still extrl...·l11d~-" low) is gellL'ratcd in
Adsorbed
boundary
lubricant
Pressutized
lIuid
~
-o~m'x;\;~,'~;
Boundary
lubricated
asperity
contact
I
Articular
surface
Schematic depiction of mixed lubrication operating in ar·
ticular cartilage. Boundary lubrication occurs when the
thickness of the fluid~film is on the same order as the
roughness of the bearing surfaces_ Fluid-film lubrication
takes place in areas with more widely separated surfaces.
Adapred from Armstrong, e.G" & Mo~-v. Ve. (1980), Friccion. Ivbr;(Miofl clod wear of synovial joints. In: R. Owen, 1. GoocH(!lIow,
ilnd P. Bullot/gll (Eds). Scientific Foundations of Orthopaedics
(lnd TraumMology (pp_ 223-232), London: William Heinemwnn.
•
I
I
i
boundary lubricated areas while mOSl of the load is
carricd by thc fluid,filrn (Dowson, 1966/1967, 1990),
The second mode of mixcd lubrication (boostcd
lubrication) proposcd by Walkcr ct aL (1968, 1970)
and !'vlaroudas (1966/1967) is based on the movc,
mCI1l of fluid from the gap between the approaching
arlicular surfaces into the articular cartilage (Fig.
3-19C). Spccifically, in bOOSICd lubrication, articular
surfaces arc believed to be prolected during joint
loading by the uhrallltration of the synovial fluid
through the collagcn-PG matrix. This ultrafiltration
pcrmits the solvent component of the synovial Iluid
(water and small electrolytes) lo pass into the articular t.::artilage during squeeze-film action, yielding a
conc~ntrated gel of HA protein complex that coats
and lubricates lhe bearing surfaces (Lai & fvlow,
1978). According to this theory, it becomes progressively more difficult, as the two articular surfaces
approach cnch othec for the HA macromolecules in
the synovial nuid to escape from the gap bet\veen
the bearing surfaces because they are physically loa
large (0.22-0.65 fJ.m), as shown in Figure 3-23. The
Water and sll'lall solute molecules can still escape
into the articular cartilage through the cartilage surface and/ol' laterally into the joint space al the pcI'iphc.)' of thc joint, Thcol'clicall'csults by I-Iou ct aL
(1992) pl'cdict lhat lluid cnt.)' into the cal'tilagc,
bearing surl"ace is possible, leading them to suggest
that boosted lubrication mav OCCUI: The role 01" this
HA gel in joint lubrication·' remains unclear, ho\\'-
ever, particularly in \'iew of the findings by Li
(1968), which dcmonstratcd thai purificd I-IA acts
a poor lubricanl.
To summarize, in any bearing, the effective mo
of lubrication depcnds on thc applicd loads and
the relative velocit~' (speed and direction of motio
of the bearing surfaces. Adsorption of the synov
fluid glycoprotein, lubricin, to articular surfac
seems to be most important under severe loadi
conditions, that is, contact surfaces with high load
low relative speeds, and long duration. Under the
conditions, as the surfaces arc pressed together, t
boundary lubricant monolayers interact to preve
direct contact between the articular Stll·races. Co
versely, Ouid-mm lubrication operates under less
vere conditions. when loads arc low and/or oscill
in magnitude and when the conlacting surfaces a
moving at high relative speeds. In light of the vari
demands on diarthrodial joints during normal fun
tion, it is unlikely that only a single mode of lub
calion exists. As )/ct. it is ,impossible lO slate de
nitclv under which conditions a particu
lubrication mechanism may operale. Neverthele
using the human hip as an example, some gene
statements are possible.
L Elastohyclroclynamic lluid-films of both the
sliding (hydrodynamic) and the squeeze typ
probably play an important role in lubricati
the joint. During the swing phase of walking
when loads on the joint are minimaL a subslantiallayer of synovial lluid-film is probab
maintained. After the first peak force. at hee
strike, a supply of nuid lubl"icanl is generate
by articular cartilage. Howcvcl: this nuid-filn
thickness will begin to decrease under the
high load of stance phase; as a result.
squeeze-film action occurs. The second peak
force during the walking cycle, just before t
toe leaves the ground, occurs \vhcn the joint
is swinging in the opposite direction. Thus,
is possiblc that a fresh supply of lluicl,film
could be generated at toc-orr. thereby provid
ing the lubricant during the next swing phas
2. With high loads and low spccds of "dativc
motion, such as during standing. the fluidfilm will decrease in lhickness as lhe fluid is
squeezed olll from between thc surfaces (flu
film), Under these conditions, the fluid exud
from the compressed articular cartilage cou
become the main cO!1tributor to the lubrica
ing film.
3. Under extreme loading conditions, such as
during an extended period of standing follow
:~
~'.''.'._--~~"~,-,.,,"'''',!,,,_,_,_----~,.....= ,.,.,,"'"<" "- " ', ,",'!'='-' ' r-'- - - - -~"=,., , ,.7' ' '"' ' ' ,) :~ , "'"r
Solute and
small particle
flow
I
T-=-0.02-1 urn
~
-=. ,., -
Hi
Articular
surface
1-=-- "-
-- -- ' -- --
..
- -11- - -1'1- --11-- - - H- - -1/- - -\/771J/ 77<ttl 777:JJ/ 7 7'ty 7r:Jj/ 7:JJ? 7
7
I
7
7 7
SUrf;:iCC
ilmlEL
rI
;'cW) :i:r'
ri,;;Cfor-r!O:C:CL:ic::-',
""
Ultrafiltration of the synovial fluid into a highly viscous gel.
As the articular surfaces come together. the small solute molecules escape into the articular cartilage and into the lateral
joint space, leaving the large HA macromolecules that, be-
edU"\:, of their' "ile', afe \H1i]bir:, tu C'~(ilP(" TlH.'se Hi, nhH:rc"
rnoluui(,s forrn d ((mu:,n\r,',ted
~;e!
ic'"" th"Hl i l,r11 thick
ti'
lubricate's the "nicul.Jr' surfi'l(OS, This hypothc,,,iled IL;bric,:;t
moclc' is
lC'ln:(~d
"i)()o51cci lubriCdtion "
~
ing impact, the fluid-film Ina)' be clilninatcd,
allo\ving sllrface~to-sllrracecontact, The surfaces, however, will probabl.\' still be protected, either b.\' a thin la)'Cl" of ultrafiltratccl
s.. . ·novial fluid gel (boosted lubrication) or by"
the adsorbed lubricin monolayer (boundmy
lubrication).
:\III](lll~~lt
i\J:li.,!
i:-- p~lrlili(IIl!..:d
h,:l\\l'l'll llll' "o
:llH,lllllid pll~l:-'l'" o! ~\ hipll~l"il' Ill~\llTi~d (\10\\ l'l ~
jl.lSOJ, .'\ll',,!JI~lll 1]l)l)7) lkri\l'd ~llll'\Pl\'""i(l!l lor t
l'lkl'li\\: ~or 1\1l'~\"\1I'l'd) l'(ll'llll'il'lll
\\~l" lkpl'lldl'lll :-.okh
(II'
friction th
lo
h\ thl' :--olid In~ltl'i\ (c',~ .. lhl' dilll'I\:IlC
hl'I\\l'l'll tOI~tl 1(l~\l1 ;\11(.1 tll;\l "lIpj1ol'll'd h\ h.\dl'o:-:
OIl
thl' proportioll or 1Ill'
:-'lIpplll'1l'd
Ii . . ' j11\',":-;[II\' ill llll' l]lIidl. Till' ilnplil';llion 01 "lIch
ROLE OF INTERSTITIAL FLUID PRESSURIZATION
IN JOINT LUBRICATION
During joint articulation, loads transmitted across
a jointma.\' be supported by the opposing joint surfaces via solid-to-solid contact, through a fluidfilm la,\'er, or b.\' a mixturc of both, Although rIuidfilm lubrication is achievable, its contribution
to joint lubrication is transient, a C()J1sequcnce of
the rapid dissipation of the fluid-filrn thickness
by joint loads, Witb tbis caveat, Atesbian (1997),
adopting the theoretical framc\vork of the biphasic
theory (Mow et aL, 1980), proposed a mathematical formulation of a boundary friction model of
articular cartilage to describe the underlying
mechanism behind diarthrodial joint lubrication,
in particular, the time-dependcnce of the friction
coefficient for cartilage reported during creep
and stress-relaxation experiments (1\'lalcolm, 1976;
McCutchen, 1962),
l'\j1I'l'",,,i(lll i" lilal Ihl' !I'il'li(l]l;tl !1),Opl'l'til'" (d c\n
L\~:'-l'
\';1]'\
\\ilh lillll' dlll'ill~: ;\ppli,'d 1(1;ldill~, ~l 1\'l
1101l ollhl' IllllT"llli~d l]lIid ;llld \.'(dl;l~:l'll-P(; I1I~l
illll'I'~\l'll(lIl" 111;\1 ~i\\' ri"l' to thl' 1]0\\ "lkpl'lldl'111 \
<..'(ll,Lt:--lic Pl'(lPlTlil'"
01
111l' li","lIl' lk:-,crilwd
\..';\r
~\Ild "ll(l\\ll in l:i~',Ul\'" ,;_l) ;llld
3" 1O.
To \'alidall' Ili:-, l1lodl'!. ;\ll':-,lli~\ll dl'\'l'lopl'd ~\ IHl
IU;ldill,~ l'\j1o,.'l'illll'lll lkll "lllkTilllpU:-'l'd ~l I'ri....-tl(j
lorqlll' 10~ld 011;\ ""';ll'til;\~:'-l' l'\pLllli 111H,hT~~oill~:'- l'J'l
1();HJin~, ill ;1 Ulillill,'d l'()lllp)'l'S"ioll l.'ollligl1r~\
(Fig, .~-~-L\l ("\tl',,hi~lll l'l ;11, ll)l)Sl. \lnrl' Spl'l
L'all.\,
;1
\.'\lilldril';J1 hip!J;l."il' l·;ll'lil~l!.-',l' plug \\'~l:-' l'O
pl'l':-""l'd ill ;\ l'()lJ1illing rillg k,~2.., prohihiling r~ld
lllolion ;\lll! nuil! l'\ud~\tion) 1l1ll!lT ~\ l'onstalll a
plil'd IO;ld ~_',l'lllT~lll'd h\ ~lfl illlPl'ITlll'~\hk' rigid pLl
llwl \\~l" r()lalill~,:'- ;11 ;\ prl':-:;\.'I'illl'l! ~lll!--,-lllar :-'jk'l'd,
T
:-'lIrr~\l>l' or thl' plllg oppO:-,itl' till' pl~\ll'll \\~l~ prl's
;lg;\ilbl ;\ li\l'd ri~2.id PO]'OU:-' IIIll'\' \\lllTl'h.\ tIll' int
di::,-it~\ti(lll or
tll~,'
Ihe l';\)'liL\!2.l' \\'lll1 lhl' rOIl!..'.ll :-'lll'!;\i.X
POI'lI\!:-' liltL'!'
prl:\'l:l~tcd it
I'n)lll
n)t:ltill~!:.
III l
,
w
Confining
chamber - - / - -
7+H---+-- Impermeable
rigid platen
Cartilage
Porous
filter
--/-----11--
--l---+':~:',,-:~.:.; :::'::l~::..\":.~:~~~.;:
I•
i
A
~
Fluid exudaricn
0.15
I
0.13
0.10
~ 0.08
0.05
0.03
BOO
B
1200
1600
2000
2400
Time(s)
Experimental configuration superimposing a frictional torque with creep loading of an articular cartilage explant in confined compression (Ateshian et aI., 1998). A. Note that fluid
exudation occurs on the opposite face of the tissue exposed to the frictional load, indicating that the frictional properties of cartilage are not dependent on the weeping of interstitial fluid to the lubricating boundary. B, Note that effective friction coefficient <.... r!')
varies with increasing proportion of load on the solid matrix, as can be seen from the theoretical curve for V-rll as a function of time during the experiment. Adapred from Mow, VC
& Areshian, G.A (1997). Lubrication and wear of diarthrodia! joints. In 11.e. l'.;low g. We. Hayes
(Eds.), Basic Biomect1anics (2nd ed., pp. 275-315). Philadelphia: Lippincott-Raven
manner, a frictional torque was developed in the lis~
Sue. Because the application or a torque load that
yields pure shear, under inflnitc~imal deformations,
induces no volume change in the tissue or associated nuid exudation, the load generated by the frictional torque is independent of the biphasic creep
behaviOl" of the tissue. Theoretical predictions,
which closely match expel"imcntal results, show
that during initial loading, when interstitial pressurization is high, the friction coefficient' can be
very low (Fig. 3-2413). As creep equilibrium is
reached and the load is transferred to the solid matrix, the friction cocrficient becomes high (e.g.
0.15). The time constant for [his transient response
is in excellent agreement with observed experimental results (Malcolm, 1976; J\l1cCutchcn, 1962), Another imponant resull or this work is that nuid pressurization can function in joint lubrication without
concomitant fluid exudation to the lubricating
boundary as is proposed for weeping lubrication
(McCutchen, 1962) (Fig. 3-190). Equally significant,
this lubrication theory is capable of explaining the
obscl\'cd decrease of the effective friction coefficient with increasing rolling and sliding joint velocities and with increasing joint load (Linn, 1968).
Recently. the inlcl'Stilial fiuid pressurization
within cm·tilage during uniaxial creep and stress relaxation experiments was sliccessfully measured
(Soltz & Aleshian. 1998). As predicted by the biphasic theOly, they found that interstitial Ouiet pressurization supported more that 90'/0 of the load for several hundred seconds following loading in confined
compression (Ateshian & Wang. 1995). The close
agreement
their measurements with biphasic theoretical predictions represents n rnajor advancement in the understanding of diarthrodial joint lubrication and provides compelling evidence for the
role of interstitial Ollid pressurization as a fundamental mechanism underlying the load-bearing capacity in cartilage. It is emphasized that while the
collagen-PG matrix is subjected to hydrostatic pres~
sure in the surrounding interstitial Ouid, it does not
expose the solid matrix (nor encased chondrocytcs)
to deformation, presumably causing no mechanical
damage.
or
Iv,:l '1111~h'\"II""lld;k't" \11111;11,:1 ])vl\\l'l'll til",
til"~ (d I Ill' 1\\1) l'~II'lil;l~l' '111'1;ll'I.."
of Articular Cartilage
'Ncar is the unwanted removal of material from
solid surfaces by mechanical action. There are two
components of wear: interfacial wear resulting from
Lhe interaction of bearing surfaces and fatigue wear
resulLing from bearing deformation under load.
lnterfacial wear occurs when bearing surfaces
come into direct conLact with no lubricanL film
(boundary 01' Iluid) separating them. This type or
wear can Lake place in either of two ways: adhesion
or abrasion. Adhesive wear arises when, as the bearings come into contact, surface fragments adhere to
each other and are torn off from the surface during
sliding. Abrasive wear, conversely, occurs when a
soft material is scraped by a harder one; the harder
material can be either an opposing bearing or loose
particles bc(ween the bearings. The low rates or interracial wear observed in articular cartilage tested
in vitro (Lipshitz & Glimcher, 1979) suggest that di-
\\
fli dkt,.·lih· l
rukd (Jill. Tltl..· lI11tllipk· 11111<,11.·,. . .
1Il!ll \\l!llill!~ ill ,,'(I1II..'l·rl ;ll'l' Illl' 1111'I..,ll:llli'JII'
t
\'"1.::11' pi :lrlit"IILII' ,::llliLI.!-'-\." IlllJik ... ·k
iIlIVr!;ll·i:d
<.·nl1<.'k- ........... ;ldh<.",j\l· ;lIld :d)r~l..... i\<.· \\l·:lr 111:,\
\:11"1
ill :111 illqxtirl·d Ill" . . k·~! . . ·It",·r:ll<."d '.\llll\i:d jllill\
[)h.' l·;II'liLI.!-'-l· :-'lll'I;ll'l' 'll,t:lill' 11111';I'll'll\.:lllJ':tl
:llHI'll!" dl'I..Tl·:I'V' ill
111:1".
il
hl·;,,:Ollh."' :-'Idi
:11. !lJSh: :\nn"'l
\1t,I.I.. 19~~: Sl·tloll 1:1 :d. I(jl..)-ll. Thu,.... lItJid Irc
111ur<.' rk·nlJl·;lhk· L\ki/ilki 1,:1
Illhl'il';llll !dlll 'l'p:lr;llil'I~: dll' hl';lrill~~ ,111'I:ll'
Il,__ ;d~
Tlli~
!\\ .... ·I..'ll
'
hd\\l'l'l1 11
lil .... i.:anil:lgc
11lhri...:;Llill~ Illiid ll'IHII
10:-.... 01
f;li.x,
pl"!
IhrClll~1i
:1\\;1\ I11UI\.' L·:l ... ih
I hI..· prllhahilil\' III dil\·t,.·1 i.-"Clll
inlTl·;l'\.'"
11lL' ;1:-'PlTilil"
;llld V\;ll'I'J"h;lk, tIll' ;lh
't,.·l'~:-'.
Ll\i~\II..' \\\,';\1"
hl·;ll·;ll.;! '1lI'Lt<.·l·~ I"l· ... IIII,... Ill
ld
,"'UI"I~ll"... :-\l'·:-'llrl~iI.,."L" ,.:lllll:tci !lUI Ii-tllli 1111..· ;H,:L"
ti('ll (d l1'li,Tl""'i..'iJpL,: d;ll)I;I.:,~L· \\ililin tIll: hV;II'il
kri:d until·' l"l·p,:lili\L· ... (rL· ......... ill~. Bl·;lrill;;: ,
L.illlrl' ,ll;t-'
11 ... '\.'111·
\\illl Illl" rl'pl":t1nl ,lpplil':l
lli~1i l('~ld ... O\er;l !"c!;tli\l"h :-hl.n lhTi(ld (II' \\t
1\'PI..'1 j I illil
lh(lll~h
1IJ\\l..'l'
I
d' It
lilt·
j( );Id ... (1\l'1" ;111 l'\ I \.'lldl"d j1l"l'il
1\\
11l~1[.!llillllk·
til'
1!J;11l !lJl" lll:lll"l"i;d":-.
thll ... I..' 1(,;((../ ... l1la\ Ik·
lilt i1ll;11l" .... 11\·11,;2111. l
lii~tll· \\l':11". r...:' .... l!ltill~ il"ll!lll..·\l·lil':llh I"l·Pl';\II.·d
Ill:lli(1I1 nldll· hl':lrill~ 111;1ll'1'i;t! ..... L';ll1 1:1 hi..' pl:tl
ill
\\cll·lllbril..":lIl'd lk·;ll'ill~:-'.
III :-\lll,\'j~d
joinl
Wear
I~ill..·h lll..'l..lll'
"j\t,.. \\l·;lr ill dh,,"V ","\pl..·rillh:l1h, 111'\\l'\<.·I".
Ill~ld
jllilH ..... IhL· ..:\I..'lil.·~d \:Iri:llillll i
dlll"ill~"
l)h\,i(ll()~il·:t1
11111,t
:ll
...':111.-.1..':- I"l'pl'lili\l' ;lrti ...: ll!:tr \:;:It"lil;l~~l.' ~lr""~~il1~,~
1ll:1IiullJ. III :u..ldi!ltlll. dUrill~~ rU!:llioll ;tllt.! ,lid
~pl.·\.·ilil..' I"l"giflll
:\11ll (llil"
01
01 IItL· ;ll"ll,,:u!:tr :-'Url~H:l·
Ill\,· !lJ;ld ... ,d l'olll;\i,:I
;\I·l':I.
·'1111
!\'pl"
~tl"l·:-...;ill~ 11l;ll aniL'u!:tr 1\'gilliL l,(';ld~ iIIlPl):
:lrliL'uL,lr L"'lrlil~I~I..· a 1'1..' ~11PPIll"!I.."d
1ll;lIl"i,\ ~1lh.1
h.\·
In'
1111..' l'ullag
rL·",i~I;:1l1(l' gt,.·lllT~llcd
IIll·
Ill()\·l·IIII..'nl t1tl"(Jll~lllllll [Ili..' Ill;ll!"i\. '1'1111"
joint
Illtl\ ...'IlIL"111 ;IIHI ll';ldill~ \\ill 1..__ ~\Il"'L'
SI1,,-',,:-.ing til" I hI..' ...;11Ii<l Ill:!l!"i\
1)
I\·p
1\'
:IIH.II'\..'IW:I1l"d ""\lI
and irllhihilioll (JI 1111..' lis:-'llL'·:-' illl\,"l".... lili;d llllid
l\: ;\tl'shiall, Il)')/l. TIII.., ...;\" P1'l)l'l'~:"l'S ~'.i\'"," Ii...;\,.'
ptl:-.:-.ihk· 1l11,(!l'llli:--llIs
h.\ \dti\:h f:Jli~ll"'· d;:lln:I~
,It,.·L·lllllll!:lll' ill ;lrlil..'ld;ll· 1..";ll'lil;l~l·: disnlpliot'!
C(JlI'l~l·II-PG ~(llid 11l:llri\ :lIld PC '\\';\~h Uti!.··
First.
!\"Pl'tili\\'
",'(luld d I:-ru pi ti'L·
l·tdl:t~~l·ll-PC; Illalri\
\."(JII;I~L·1l Ii IlLT:-..
I hl"
:-'t
PC 1ll:ll
l'l'lIk:-.. :11)(1'01' Iltl' illll'd;:Ii..·I..' hl'I\\n"l1 Ii\l·~l' 1\\1
PUllL·"t",.:\ PU\llt!:ll' Il-,polhl",is
i.. .
111.,\
li~lIL' i ..... Ihl' l"L· . ..;ltll ld';l It.'lbiJ...· I':lihll\'
lii~",'l"
111.,:1\\0,"1, (FrL'I..·lll;lll.
t,.·;lI"ti\;,
PI' I hI.' l·'
1075)" ;\Isu. :IS di:-l
~bove pronounced changes in the articular cani-
(-~e-':l;G population have been obscn'cd with age
:ric\i_lisease (Buckwalter et aI., 1985; iVluil~ 1983;
\R<Juahlev et at. 1980; Sweet et aI., 1979). These PG
';. chal~ges "could be considered as part of the aCCUl11l1idrcd,tissue damage. These molcculaJ- stnlclural
:-'~ha'nges would result in lower PG-PG irHcraction
;~ites and thus lower network strength (Mow et aI.,
1989b; Zhu et aI., 1991, 1996). Second, repetitive
:and massive exudation and imbibition of the intcr<~iitii:d fluid may cause the degradcd PGs to "wash
, ollt~~;::from the ECM, with a resultant decrease in
stff0e~s; and increase in permeability of the tissue
tha('jp::lurn defeats the stress-shielding mechanism
oUnierstitial fluid-load support and establishes a vidous/cycle of cartilage degeneration.
A third mechanism of damage and resultant articular wear is associated with s~'novial joint impact
loading-that is, the rapid application of a high
l8"ad. \,vith normal physiological loading, articular
cartilage undergoes surface compaction during the
compression with the lubricating fluid being exuded
lhi'ough t.his compacted region, as shown in Figure
3..:10::';\5 described above, however, fluid redistl'ibutiol}.',\vithin the articular cartilage occurs over time,
which relieves the stress in this compacted region.
Thi~'process of stress relaxation takes place quickly;
the stress may decrease by 630/0 within 2 to 5 seconds (Ateshian et aI., 1998; ""lo\\' et aI., 1980). If,
howevel~ loads arc supplied so quickly that there is
insufficient time for internal Iluid redistribution to
reHeve the compacted region, the high stresses prodliccd in the collagen-PG matrix may induce damage (Newberry et aI., 1997, Thompson et aI., 1991).
This' phenomenon could well explain why Radin
and Paul (1971) found dramatic articular cartilag
damage with repeated impact loads.
These mechanisms of wear and damage may b
the cause of the commonly obser'ved large range
structural defects observed in anicular canilag
(BlIliollgh & Goodfellow, 1968; Meachim & Fergi
1975) (Fig. 3-25, A-C). One such defect is the spli
ting of the cartilage surface. Ven ieal sections of ca
tilage e:\hibiting these lesions, known as fibrillatio
show that they eventually cxtend through thc fu
depth of the articular cartilage. (n other specimcn
the cartilage layer appears to be eroded rather tha
split. This erosion is known as smooth-surfaced d
structive thinning.
Considering the variety of defects noted in arti
ular cartilage, it is unlikely that a single wear mec
anism is responsible for all of them. At any give
site, the stress histor:v may be such that fatigue
[he initiating failure mechanism. At another, the l
brication conditions may be so unfavorable that i
terfacial wear dominates the progression of cart
lage failure. As yet. there is little experiment
information on the type of defect produced by an
given wear mechanism.
Once the collagen-PG matrix of cartilage is di
rupted, damage resulling froll1 any of the three we
mechanisms mentioned becomes possible: (I) fu
ther disruption of the collagen-PG matrix as a resu
of repetitive matrix stressing; (2) an increase
"washing out" of the PGs as a result of violent nu
movement and thus impairment of articular car
lage's interstitial Ollid load SUppOrl capacity; an
(3) gross alteration of the normal load carria
mechanism in cartilage, thus increasing friction
shear loading on the articular surface.
I
I
I
Photomicrographs of vertical sections through the surface of
articular cartilage showing a normal intact surface (A), an
eroded articular surface (8), and a vertical split or fibrillation
of the articular surface that will eventually extend through
the full depth of the cartilage (C). Phocomicrographs provide
chrough che courcesy of Dr. S. Akizuki. Nagano. Japan
All these processes will accelerate the rale of interfacial and fatigue wear of the already disrupted
cartilage microstructure.
Hypotheses on the Biomechanics
of Cartilage Degeneration
ROLE OF BIOMECHANICAL FACTORS
Articular cartilage has only a limited capacity for repair and regeneration, and if subjected lO an abnormal range of slress~s can quickly undergo total failure (Fig. 3-26). It has been hypothesized that failure
progression relates to the following: (1) the magnitude of the imposed stresses; (2) the total numbe,· of
sustained stress peaks; (3) 1I1C changes in the intrinsic molecular and microscopic structure of the collagen-PG Illalri.'\; and (4) the changes in the intrin-
Cartilage Structure
t
h.
Biochemical Composition
t
CaliagenlPG ECM
Mechanical....- Joint
Propenies
Loading
+
Spatial and Temporal
Slress-Slrain, Pressure
and Fluid Fields
1
ChanreYleS ·41---- Cell Slimuli
Synthetic Activities
~
Cartilage Function
Physical Activities
Flow diagram of the events mediating the structure and
function of articular cartilage, Physical activities result in
joint loads that are transmitted to the chondrocyte via the
extracellular matrix (ECM). The chondrocyte varies its cellular activities in response to the mechano-electrochemical
stimuli generated by loading of its environment. The etiology of osteoarthritis is unclear but may be traced to intrinsic changes to the chondrocyte or to an altered ECM (e.g"
resulting from injury or gradual wear) that leads to abnormal chondrocyte stimuli and cell activities.
~i<: IlK'chanical prop<..Tt~· 01"
the tissue. Tht: nH)SI im
ponam I"ailure-inilialillg rac((1l' appears I(J he I
"loost..'ning:" of tile c()lla~cn nctwork tllal al!c,ws a
normal PC c.\pansiclIl and thus tissue :-,weJ[i
(:\ilaroudas, 1976; McDL'vilt &: Muir, 1976). Asso
alL'd willl Ihis Chi.U1gc is a dL'Cn:i.ISL' in canibg.1.: stil
Ill:SS and an incrl.:ase in canilage pi..'rI1lL'abil
(Allman el aI., 19S-l; Armstrong &. Mow, 198
Guilak L"l aI., 1994. SL"lton el aI., 1994). h,,'h
which aller canilagL' fUllction ill ~\ diunhrodialjoi
during joilll mOlion, as showll in Figure 3-27 (!vlo
'" Atcshian, 1997).
ThL' magnitude of the stress sustained b,\' lhe a
ticular cartilage is dctennined b,\' both the toti.lllo
on the joint and how tlwt kli.\d is distributed ()\
the i.\1·ticul~ir surface conl~lCt arca (Allmed 6.: Burk
19i'3; Annslrong L'I aI., 1979; Paul. 1976). :\n.\·
h,:nsl..' stress COllcL'nlration in Ilk' l'ont~\Ct i.lrea w
pla,\' i.1 pl·imi.llY roIL, in tissu<..' (k'gL'Ilt.'ration, A lar
Ilumb<..'r or well-knowll condilions GillSi...' L'.\cL'ss
stress conccnlrations in anicular canilage alld
sult in caniJage failure. Most of Illese stress L'once
II':Hiolls are C~ltISI,:d h.\· joint :,urfacL' irH..' ongruit.\',
sulting in i.lll i.\hnornwtly small conti.lCt are
E.xi.lmplcs or conditions causing stich joint inco
gruilil.:s include 0:\ subscqllL'fl( to congeniwl a
L"!i.lbular dysplasia, a slipped cnpilal femoral <.:pi
.,"sis. and intra-articular fractures. Two fUrih
eXi.\mpk-s arL' knet: joint l1lcnisct.:uoll1~·, \\'hich eli
inak's the load·dislributing function of the lll~n
ellS ([vlow Cl aI., 1992), and ligi.llllCnl I"llpllll"<"', whi
allows cxccssi\·<..' mo\"t:'1lll'1lI and the gen<..'r;:llion
abnormal mechanical ~trLSSl:S ill Ihe afk'clI:d jo
(Allman el aI., 1984; Guilak Cl ai, 1994; ivkDe\"ilt
Mui,·. 1976; Sellon el al.. 1994). In all the ,11)O
cases, abnorllwl joint ~\rtj(-ulation increases t
slJ'i..'SS acting 011 tilL' joint slIrf~lce, which aPlk'ars
predispose Ihe caniiagL' to railllr<..'.
iVIHcroscopiG\II~', stress IOGlJiz<.ltion and cOllce
tration at the joint surfaces li~I\'L' i.\ I"url!ler dlt.'C
High COntact pressures bctW('L'n Ihe anicular su
faces decrease Ill<..' pr()babilit~· of fluid-fIlm lubric
tion (Mow &. Aleshian, 1997). Subscqllcnl <.lelL
surl'''lI.:e-lo-surracc contacl
i.\Slx.'rities will cau
microscopic stress concentrations that arc rt.'SI)(>
ble for further tissue damage (Atcshian et aI., 199
1998; !\leshian '" Wan!'. 1995) (Case Study 3-1).
The high incidencL" or speciiic joint dcgerlt'rati
in indiddllals with certain occupmions. such
football pl<.l~·L'rs' knccs and ballet danccrs' i.lnkl
can bL' <,,'xplainL'd b\ tilL' incn:i.lsL' in high i.lI1d i.lbn
Illi.llioad rn:.·qllcnt:~· and magnitude sllslainc:d b~' t
joinls
Ihese: individuals. II has becn suggL'slt
or
or
•I Deformation
,
fluid
QA Progression
Normal
Cartilage
..--.
I
exudal.on ~
PG Loss
ColJagen
Damage
I
Load
. i.~yll.
: ~ -::t_~+f
(e.g., inlerleukin-I) (Ralcliffe el aI., 1986) a
growth Factors (e.g .. transforming growth fact
bela I) also appear lO pia\' an importanl role in O
Another contributing factor to the etiology of
ma~' be age-related changes to the chondroc
(Case Study 3-2).
I I I ,--'--
I
I
I
I
I
., _.1 _'--_ _--'
IMPLICATIONS ON CHONDROCYTE FUNCTION
The ECM modulates the transmission
or joint lo
to the chondroc,yte, acting as a transducer that c
1Fixed Charge Density - - ! Swelling Pressure
! Frictional Drag - t Hydraulic Permeability
verts mechanical loading to a plethora or envir
mental cues that mediate chondrocyte function.
healthy articular cartilage. loads from norm
joint function motion result in the generation
Increased Matrix Deformarion
Increased Fluid Flow
Acts 10 diminish cartilage load-bearing properties.
Knee Meniscectomy
0rty-year-Old man who had a meniscecwmy 10 years
figure illustrating how osteoarthritic changes to the col'aqen·PG network can compromise the ability of articular
ca'til'lOe to maintain interstitial fluid pressurization, which
""'~"'''o< the tissue's load-bearing and joint lubrication ca·
Loss of PG and damage to the collagen fibers result
an increased hydraulic permeability (decreased resis-
!'c!"""e to fluid flow) and supra-normal loads and strains on
solid matrix (and chondrocyte).
F
ago in his right knee. Currently, he is sufiering pain as-
sociated with movement, swelling, and lirnilaiions of knee
motion (Fig. 3-1-1),
The history of knee meniscectomy not only implies an
alteration in joint surface congruence but also the elimi·
nation of the load-distribution function of the meniscus.
The effect is an abnormal joint. characterized by an increase in the stress acting on the joint surface that results
in cartilage failure. Most of these stress concentralions are
;~ ~.;,f'
thiLin some cases, OA mav be caused by deficicn'<eie'~3~' the mechanisms th~t act to Il1ini~llizc peak
'~'ff6rc,~s~on the joints. Examples of these mechanisms
~jh~~l~(le the ~clive processes of joint nexion and
_~:~,q~e, lengthening and the passive absol"ption of
':;k,,,by the subehondral bone (Radin, 1976) and
ciscus (Mow el al" 1992),
~generativc changes to the structure and com~gion of articular cartilage could lead to abnorkti:ssue swelling and funZtionally inferior biome~
;;,iji8<11 properties. In this weakened state, the
'~-,:J~ge ultrastruclure will then be gradually de.-g.t~d by stresses of normal joinl articulalion (Fig.
~m)j7;OAmay also arise sec~ndarily from insult ~o
:' ':- ~~i~tIinsic molecular and microscopic structure of
/;f.~tfi~2so;qagen-PG matrix. Manv conditions may pro:\-.Iil<?!~:~,l!ch a brea~down in .~latrix .integrity;· the~e
-;111~e degeneratIon assoclatcd wllh rheumalOId
--·.1isJ joint space hemorrhage associated with
'p.pilia. various collagen metabolism disorders.
}~S§.tl.e degradation by proteolytic enzymes. The
_,'l1C: c ': of soluble mediators such as cytokincs
caused byjoint surface incongruity, resulting in an abnormally sma!tconraet area. This small contact area will suffer hign.c9_ntact pressure. decreasing the probability of
fluid-film lubrication, and thus the actual surface~lo~
surface contact will cause microscopic stress concentrations that lead to damage.
Case Study Figure 3-1-1,
~------_
.._-----~
Osteoarthritis
eventy.year-OJd woman, overweight, with OA of the
S
right hip joint with associated symptoms of pain, limi-
tation of motion. joint deformity, and abnormal gait (Fig.
3-2-1).
OA is characterized by erosive cartilage lesions, carti-
lage loss and destruction, subchondral bone sclerosis and
cysts, and large osteophyte formation at the margins of
tho joint (Mow & Ratcliff., t997). In this case.
roentgenograms of the right hip of the patient show a
decrease in the interarticular space and changes in bone
surfaces as sclerolic and osteophyte formations.
nle most
severe alterations are found at the point of maximum
pressure against the opposing cartil(1ge surface, in this
case at the superior aspect of the femoral head.
... t1I........ 10\\ j1l..'l'Illt.,<thilit.\·, thl..· n(lnll~tI em il"UIlIIll..·IH ell
illl..· ChllJJdrlil:.\·!I..· i df)lIlill~lil..'d h.\· h.nlrli:-'I~ltic 1'1'1.......:-'111"1..' ill the illtL·r tili~d lluid. \'~tl·i(jth phl..'lI(lllIL·II~
origil1;llitl2:-' IrOll1 inIL'I"slili~t111uid Ilo\\' l',\ist ~IS \\l·11.
llJ)plil'~\kd ill I..'nh;lIll..·ill~ Illitril...'llt dilfliSioll, illlL'!"
:-.{iti~d lIuid 11<1\\- (i.I.,:..
unhoulld \\'~tlL'rJ ~i\'L'~
ri"L' 10 l'1..'J111I~\1' ....lillHlii Ill' ~lll 1..·ll..'cll'k,d 1l~lllll"L',
n~lnwl.\' ... 1l"l'~Hllillg p(ItL"nti~d~ and (,.·UITL"JllS {Fr~lllk
Gr()t!l.insk\-. 11.)~7: Gil L't ~J1. 190.). 1t)t)t'). III ,Iddi~
lioll, illk'rslili~d lluid 110\\' tlJr()lI~h tilL' "mi.lll p0l"es
a.... :-.(J(..·i~·lk·d with tilt.' ....olid Il1;'HI·j\ (-?iO IlIll) or Ilol"lna
\:al'lila~l..', \\hkh olll'r 1..·(Ilbi<.IL-rahk rl..·si:-'l~lIKl· ti
Jluid 111)\\· (:\lanlud;'I:--, I ':lit): ,\kCtltchL'Il, 1<)6]: \tluw
1...[ ~d .. ll..)~..j.). \:~1I1 2:-'i\'L" l"iSL' [0 ~\ Jlll'l'h~\llk;l1 phl'Jl()1I1
1,.'lla tl..'l"lllL'd Iluid-indlll'l'd 1\1~\Il"i.\ compadion (Lli I.\.
\-10\\·, 1'J~OI. ThL' frktiollal illk'l"~lcli(Jn hL"l\\-t.·\,.'1l inkl"slili,d Illiid and sldid ~\rl' <l l"L·....ull
drag l"L'"ist~lllL·L· to r(JI·I..·I..·d i1U\\ Ihrullgh Ihl.." pllroll . . ·lkTlIll..·ahk
",::ll"lila~l' 1I1~\1I"i,\ ~\Ild ~l \'iSCllll." sl'IL'~Il" sll'l'ss L'.\l·l·ll'd
hy IhL' intl..'r... tiliailluid. Gi\'(,.·ll tlll" nominal !lo\\" l'aks
01 lhl.' ilHt.'l":-.tilial i111id llh..·IUiulll.."d I..·adil..'r ~1IH.1 Ih.
10\\· pl·rrl1l..'~\hilil\· "I lIll..' l';ll"lil;qll' Irl'lll'i\. cholldruC.\ II..' pl'l'ct.·ptilll1 of thi .... frictional illh-'ra(,.·,itlll f(Jl"i..·l· is
likL'h to Ik' dUlninalL'd \1\' thl..' dl"~\g rL'si:'oI~IIlCL'
!lO\\
Ihl"()lJ~h [III..' IlJ~\lri\ l"allll'I' [ll~\Jl h.\ diJ'l·\.."[ \,i:,c()\.IS
~hl..'al· :'trL'S~ l,Jl I hI..' cL·II. This friL"tjonal dr~lg forc(.
1..:<111 pr<'(\UL''''' ....u!id 11\;,\[,.;\ I..krOnll~t(i(1I1 011 1111..' urde
(II' 15 10 .'(V-"
Frum tilL' disctls:-.ioll ahO\'L', i..'l'lolldI'Ol·.\-IL' (1...'1'01"
m~I[i()Jl L>~\1l hI..' UlJJ~i<.lI,'I"L·d 10 hI..' gO\·lTI1L·d hy [hrL'e
cuupkd loadjn~ Illl..'l'hallislll~: dirl..'ct I-:C.\;, <.IL-for
tl1~llioll: Ho\\··;ndllL·l·d L'(ll1lp~\ctiotl: and fluid prl'S*
sllri/'llilln. III OJ\. I hI..' iIlI..T(,.·asl..·d lis~lll..· pt.·nlll·'lhil~
il.\" diminis!k· ... carlibgL"~ !lormal flu;d pr..."~~lIre
Itl~I<.I·SllPP()I"[ IllL'L·l1ani."lll. Thus, [i'lL'!"...' i." a shirt
or !O,ld support Ollto thl..' solid matrix. I..·i.lusing
sllprallol"nwl ,slrl..·S::>L·:' and s!l·ailb 10 hI.." illlposl.'d
011 1111..' chondniL·.\·!I..'S (Fig. 3-271. Th..:·si.' ahnormally high Stl'L'SS ~lnd . . tl'~lill (('\-C Is , and utili..'!
rn(",,,,'k\llo·l'iL'I.:II'Ol"Jh'llli(,.·~\ll·h~lll!:!I...'S [Iw[ ~\rL" nwnik~l('d \\'itll 0;\, C=ill triggl..·1' an irl1hal~IIKI..' of <.:IHIll
<',,"01..'.\'1\.' an;:lholic and I..·alaholic al'li\'jtk"s, further
l,.'Ollll"ibuling In ~l \'icious c.n·1t.- Ill' progr...'s~i\"i.: Glr!ilagl" dl.."gl.."nl..·r~lli(lll. Itllk...' d, changes to tht.' bin
L'hL'lnkal l:Olllpositioll ~llld s[l"lIclUI"l,: oj' cal'til~\g
Ctlll Iwu..· a profound inlp~lcl Oil tiSSlll..· and (,.'!lOIl
d n)cy{t.' fUlll'1 ion. \\'i t h IntI! t iLl i:,dpli nary collaho*
r~ltioll:' =md ~1Il appropri::lli..' thl·Ol"I..'lil"=-d rl'~IIllC*
\\,(H-k, such as th ...' biphas:ic tl1 ....<I1·.\·, iJlsight~ illW
'Ii'lL' factors that g()\·1..'1'1l c!londnlC\[<.' f"llIKlion
L'arlil<lgL' s[ruclllr...' <lnd l'UI'H,:lioll, ~\IH.l lhc diology
or 0:-\ I..·an hI..' obtain..:d,
or
«
or
or
Case Study Figure 3·2·1.
mcchano-e1ectrochemical stimuli (e.g,. hydrostatic
pressure, SlI-ess and strain fields, streaming potenti~t1s) that promote normal caI·lilage maintenance
(by the chondroeytes) and normal tissue function
(Fig, 3-26). However, when the integrity of the collagen-PC network (the transducer) of articular cm'tilage is compromised. such as from trauma or disease. normal joinl articulation leads to abnonllal
Illcchano-clcctrochemical stimuli, with ensuing abnormal ECM remodeling by the chondrocytes and
debilitated tissue function.
In the absence of joint loading, the normal environment of the chondrocyte is characterized by the
pre-stress established by the balance between tension in the collagen flbers and the Donnan osmotic
pressure. During joint loading. by vit·tue or the lis-
,
~.,t,--.
....'.~
., .. "
REFERENCES
SumrlzalY
The function of anicular cartilage in clianhrojoints is to increase the area of load distribution
;'.(thereby reducing the stress) and provide a smooth,
(\,vear-resistant bearing surface.
i;~(2) Biomechanically, articular cartilage should
!J~~viewed as a multiphasic material. In tcrrns or a
}biphasic material. articular cartilage is comprised
,i' porous-permeable collagen-PG solid matrix
(approximately 25% by wei weight) filled by the
fl'e'~ly, movable interstitial fluid (approximately
i5r~',:bY wet weight). In addition to solid and Fluid
:tIlel,~e', (~xists an additional ion phase when CO!1si~f~r'ing articular cartilage as a triphasic medium.
/;:The ion phase is nccessa'"y to describe the swcl>~cling and other electromechanical behaviors of the
.of
:'~~/tissue.
'lH~:-'{~' lmporlanl biomcchanical properties of articular
;:',c~rlilage are
the intrinsic material properties of the
soli~1 matri:\ and the frictional "csistance to the flow
9f/':,~nterstitial fluid through the porous-permeable
:spHcl,';',JTlatrix (n parameter inversely proportional to
,~h8i+issue permeability). Together, -these parameters
,;~;~,~efine the level of interstitial fluid pressurization, a
:"'-};£6ajor detenninant of the load-bearing and lubrica/iZ.J!ibn capacity of the tissue, which can be generated in
)s:cartilage.
'~,ti!i-~t;Damage to articular cartilage, from whatcver
/~%g~use. can disrupt the normal interstitial fluid loadf;"9~~ring capacity of the tissue and thus the normal
"IH§n,fation process operating within the joint.
~n~tTf()re, lubrication insufficiency may be a pd:m~~;y':(aclor in the etiology of OA.
/i?~@;,-
:.'
;"G~hr~J'Nhen
describing articular cartilage in the conframework such as the
J~.Jjlphasic. triphasic. or multiphasic theories. it is
':ltnqssible 10 accurately predict Ihe biomechanical be?:~}i~viors of anicular cartilage under loading and to
W~1,~!:;lJ~idate the underlying mechanisms that govern
';r1M,01oad-bearing and lubrication function. Further'~?,r~',:')nsights into the temporal and spatial nature
'?Rr.\p·?'phy;ical stimuli that may affect chondrocyte
·;flinction in situ can be gained.
,:D'Jexl ()f a rigorous theOl-etical
,.
(~j'r!;:
-
4?£1.CKNOWLEDGMENTS
(~~'i::;';"-'
-f~~O;:his'wark \vas sponsored bv the
atianal [nstitutes
}iitiR+lealth grants AR41913 ,{nd AR42850.
,:~:<ii~BX~(/
~. ;;,::/';:;<,-,
!\hllh:d, i\"\1.. (\: Burkt.:, D.L. (1983). In vilro mcnSliremen
st:llil: pressure distribution in synovial joints-P,lrl 1:
i:1I surface of the knee. J Biolllech Ell.!!,. 105, 216.
Akizuki. 5., \lo\\'.
"-,lulieI'. F.. (.'t a1. (1986). Tensih.: p
"nies of knel." joint cartilage: 1. Influence of ionic: co
tion. wcighl bearing. and fibrillalion on the tCllsill." mo
IllS. J Orr/lOp U/.:.\', 4. 379.
Ahm~ln, R.D., Tt:nl'baum, J., Lilla, L. l'l "I. (1984). Bio
Ch~1I1ic.:l1 <lnd bioc:ht:mical propenh:s of dog cartilage in
peri mentally induced oSlcoarthritis. All/I Rhl!/lI11 Dis. 43,
:\ndri~u.:chi, T.P.. Natar~lj'tn, R.N., ~ Hurwitz. D.E. (19
Musculoskclc·t~d dyn,llllics, locomolion, anti dink'll ~l
l.:;llioll. In V.C. Mo\\' & \V.e. H<'lyl'S (Eds.). Bw;;c
I!wpw:dic BiViIlt.'dlllllics (2nd l."d.) (pp. 31-68). Phila
phia: Lippincott-RHvcn.
Armstrong, CG .. Balmllli, A.S., ~\: Bardner. 0.1.. (1979)
\'itl'o measurement of articular c;lrlilage deformation
lh~ intacl human hip joinl und"r load. J B(m" loi/a S
61,1.744.
:\rmslrong. CG., & Mow, V_e. (J980). Friclion, lubricil1ion
w,,'ar or syno\'i,al join Is, III Owen, R.. Goodf...!low, J. (\:
lough, P. (Eds.). Sci~'IlIiric FOlll1dt/tions or OrtJ/(JI'l/~'tlic:s
7iwf/lllUO/Ogy (pp 213-232). London: William 1-ldncnn<'H
ArmSIr<)IH!. e.G., & l\.fo\\'. V.e. (1982). Varialions in dll."
Irinsic-mcchank<.tl propt:r1ies of hum'lll anil:ulrtr ci.lni
Wilh agL', degencr;ltion. and w:lI..:'r COrllenl. } 8(111~' J
v.e.,
Sltff::, 64..1, 88.
Ateshian, G.A. (1997'). Theoretical formulation for bound
rrktion in articular cartilage. J !3iOll/cc), ElIg, 119, Sl.
:\LL'shi~tn, G.A., K\\'~tk, S.D.. Soslowsk.\', LJ., et al. (199n..:'w sll:n:ophotogr:lInmctry lll..:'lhod for dctermining
situ conlact are;\s in dianhrodinl joinls: A compar
sludy. J Biol1/~'dlallics, 2i, III.
Atc:shian. G.A., Lai, W.M., Zllli. \V.B., (:t a!' (1995). An :lsy
totic solution for the contact of twO biphasic c~nibgc
\.'1':;. J Biol1lccJulIIics. 27, 1301 i.
Ateshi:lll, G.A., & W'\Ilg. H. (1995). ;\ thl'oH:tical solution
til(.' friclionless rolling contact of cylindrical biphasic
ticular cartibge layers. J BiolIlCC/HlIlics, 28,1341.
Att:'shian, G.A., \Vallg, H., & Lai. W,M. (1998). Thl..' roll.' o
tcr::;titi,d fluid in pressurization ilnd surface porositie
the boundary friction of articular cartilage. 1 hi};ol
120.241.
.-\Ieshian. G.:\., Warden. W.H., Kim. J.J .. d al. (1997). Fi
("-,formation biplwsic malerial properties of bovine ,Ir
!;Ir cartilage from confined compr..:ssion experimellt
BiulI/t.'c/UlIlics. 3D, 1157.
Athan<lsioll, K.:\ .• Ros('llwasser, i\·I.P., Buckwallt.'r. J.:\., e
(1991). Inlerspecies comparison of in sill! Illcdmn
propl."rlies or diswl rcrnor,ll cartil:tgc. J art/lOp Ik.., 9,
Bachrach, N.M., VlIlhmu, We., Stazzonc, E.J., et 'II. (19
Chang~s in protcoglycan synthesis rutes of chondroc
in articular cartilage arc associ~lted with the time de
dCllt changes in lilt: lllC'chanie<d environment. 1 13i
ch(lIlic~, 28, 1561.
Bachr'Kh. j\!'~C Mow. \i.C, & Guilak. F. (1998). Incompr
ibility or the solid matrix of articular cartilage under
hydrostatic pressure. J BioJ11£'cluwics, 3/, 4015.
Bateman, J.E, Lamanue, S.R., & R,II11SIw,w, J.A.~'1. (1996).
lagen sllpcrbmily. In W.D~ Comper (Ed.). Extracellular
Irix (Vol. 2. p. 226i). :\lIlsll'rdam: Harwood Academil: P
Bollet. A.1., &: Nance. J.L. (1963). Bio<:hcmical findillg~ in
normal and oSh:oarthritic <:tnit:ubr canibgc, II: Chon·
droitin suJf~\IC concl.'lltralion and ch'lin length. and W;:lt('l'
,.\[1<1 ash content. J Clill l'II't'St, oJ). 11 iO.
Bowden, F.P.. & T~lbor. D. (1967). rfiction /llId Lllbrictlli()J/.
London: \lC1hucn Pubs.
Broom, N. D" &. Sil~·n-R()bens. H. (1990). Collagl'n·collagcn
\'crsus collagen.prolCoglyc:w interaclions in the de!l:rmin;:ltion of c.lt·til~lgc stn..'ngth . .·Irthrilis 101i.'1II11, 33, 13[2.
Buckwa!l\:'r, J.,\., KucHner, K,E., &: Thonar. E.J.-M.:\. (1985),
,\g('-rl"l<lted changl's in articuhlr c:lnilage proleoglycans:
Eleclron Illicroscopit: sllIdics. J Orr/lOt' Re:., 3, 251.
Bullough, P.G .. &: Goodf,.'!low, J. (1968). The significance of
th{' fillc structures of ~lrticllbr cartilage, J BOllI.' Join/ Surg
50B, 852.
Bullough, P.G .. &. Jag,-Illnalh, A. (1983). The lI\orphoJog~' of
the calcific'ltion front in articul:lr l.:anilagc . .I Bouc Joiut
SlIrg, 65B. 72.
G:trdlll·r. 5.1.. ...\: .\-kGjllin;ty. D.C. (1971). Li\"ing :lrticul:I
l';lrtilagl' i:- nOI SlIllll__llh. Till..' :-trth,:tl1l"l' (,t 1ll:lIll!ll;di;lll an
a\"i;lll joint :-Ilrbl.:l':- dCllHJIl:·dr:lll..'d in \'i\"o b~ inlflli:r~i(l
indd('llt light miu(Js("op,\". ..1111/ Rhl:/1l1l Ois, 30, 3.
Garg.II.G., & Swann. n.:\. (1981). :\g,:-rdatct! ch:tng,·s lllthl
chl..'lllicd l:omposition (Jf h(I\'inl' :Inil'ld~\r ,;;anilagt
Hiot.:ht'lll J. 193, ·!.';9.
Gll, \\'.Y., Lai. \\'.\1. l..\: .\11)\\", v.e. !,1993J. Tr,-lll:-pon of fILl
Buschm'llln. l\I.D .. Gluzband. Y.,\., Grodzinsky, A.1., L't :\1.
(1992). Chondrocylt'S in agnrosl' culture s~'nthcsize a Illechanic,lIly functional i.'xtracdltll;'lr m:ltrix . .I Or/hop Res.
10. 745.
Buschmann, :\-1.0.. &: Grodzinsk...... '\.J. (1995). A lIlolc~lllar
modd of protL"()gl . . . can-nssoci~\{cd ckctrostatk forcl.'s ill
l,;arti!:l~(' mechanics. } BhlllH'ch Ell!!" 117, I iO.
Clark, J.~,i.·(1985). The organization ~f coll:lgl'll in cr.. . ofr;lctun:d rabbit ~niclllar cartilage: A s'::lnning l'll'.:troll microscopy stud...... J Or/hop lk~, 3. li_
Clarke. I.e. (19il). Articldar cartilage: A re\'ie\\, and scanning
l']cctron microscope study-I. The inll'r1crritorial fibrillar
'.l!"chitccltIl'c. J BClIIl' Joim Surf;. 538. 732.
Donohue. J.M .. Buss, D.. Oegcma, T.R .. (.'t al. (1983). The dfects of indirect blunt trauma 011 adult canine articular
Gujbk. F.. R:lIcliffL' . .-\ .. L:tlll·. :\ .....·1 ;t!. (1994)..\kch:1I1ic
and bindlcmil';d l·hallgn ill lhl' Slllh·rlil.:ial ZUlll..' of ;-\rtil..·
1;11' canil:lgl' in a ~'~lnil1(' Illo(kl or nSll..'o:lrthrilis. J Orr/w
Rt's. 12, .ti-L
flardillgh:llll. T. .\1.. l..\: .\Iull', H. (197-+1. IlYllluroni( :H.'id ill ca
lilagl' and prokogl~'C:1l1 :lggn..'~~l\ion. Hiu(·/tt·//t.l. 1.39. 56S
1·lardilJ.~dlall1. T.E .. !.h:ardrn(Jrl'·G;lrt~. \1., ~ Dunklill. D.G
(llJ8iL Pn,Il'in d(lll"lain Stl"lll"llll"e of Ih ...· aggrq~~l!illg pl'O
teogl.\T;11l frolllc\rti1:lgl·. Tnm., On/t{J!, Rt'." ,)0('.11,61.
Ibsc;dl. V.C. (l<Jill. 11I11..·r~lC(i(JllS (If l'anilag...· pro(('oglycitl
with h,\';duronil' acid . .1 S/lpl"iul!l!i SII'I!ClIll"t·. 7. 101.
Ib~·l·':-. W.c.. & Bodilh', :\.1. (1978). Flo\\'-indi..'peIHknt \'i;
cOl:l:!stir propl·nil..'s of articul;l], <.:ani!:lgl' m;\lri.". J Bitt/IIt
c/ulIIics. II. 407,
1[;1.\·('S. W.C .. l..\: \Iockros, I...F. ([971 Oi.Viscoi..'!astic propl'rlii.
of human anit:ubr canil;lgl'.) :\PI'I Phy.~i(Jj, 31.562_
1I~I\\'killgs. S.\\'. (19SSI.:\ Brid l1ililor.,· of Tillie.' Frolll tilt: Bi
Ba/l~ 10 l3!ack 110/n. Xl ·\\· York: B:IIlI<lm Book::..
Hcinl'gard. 0 .. \\'it:slall(kr. J,. Shl'dlan. J., l'l ;d. (1983). S('
:tration ;'tlld . . haral·ll.'riz,llinn (If 1\\'0 poplll;ltions of ;1g.gr
g:lllng proll..'ogl~"l';ln:- from i.':lnib~l'. Bif,d/"III J. 215. 95,
Hdminl'n. H.1,. Ki\'ir:llll;l, J.. T;lI11l1li . .\1.. L'l :.1. (Eds.) (19~i
./oil// l.orldill,!.:: Biology (lild 11(.'111111 (>1" ..\I"/iclilal' Sll"/Iel((1""
Bristal, V.K: Wright &: Sons. Pubs.
lIill:-. B.A. (1989), Oligolallll'llar luhri\:alion of join\:- hy su
fan' ~H:li\'c phnsphlJlipid. J Rlll'lIlJI. I-b. S2-91.
Hirsch. C, (19-+4). TIll' p;lIhngL'!ll'sis of l'llOndronwlacia of 11
p~lli..'lb . .-\0(1 Chi)" Sl.'{/!ul. 83 lSuppl), I.
111:1\·a<:l,k. :\-1. (1995). Till' rnll' of s~--n('\'ial fluid fillr;llioll h
l·~lI"lilag ..., in luhril':llion of s~-Il(J\"i:1I join Is: IV. Squl.'l.'1....· .fil
c;\l·tibge. J BOlle JoiJl/ SlIrg. 65A. 948.
Donnan, F.G. (1924). The theory of membrane equilibria.
Ch.:mical RCl';C\\', J, 73.
Dowson, D. (1966/1967). Modes of lubricalion in human
joints. Proc IlIs{ Mt!ch ElIg, IBIJ. 45.
Dowson, D, (1990). Bio·tribology of natural and rcp!accIlH:nt
joints. In V,C. :"'Io\\', A. Ralcliffe, S-L.Y. Woo (Eds.l Biollll.'('''(/Ilks
f)ial'thmdia! .foill/S (pp. 305-345). New York:
or
Spri ng~r- Verlag.
Edwards. J. (196i), PhysiGII charach:risti(:s of <lnicubr canilage. Proi' IIl.H :\ft:dl Ellg, ISIl. 16.
Elmorc, S.~·I., Sokoloff, L., ~orris, G.. et al. (1963L Nature of
"imperfect" clasticity of anicLll'lr CarLil~lgc. J :\pplid Physiol, 18. 393.
Eyr('. D.R. (1980). Collagl'll: Molecular di\'ersity in lhe body's
prol('in scaffold. Scicl/n:, 207, 131 S.
Fos~Hlg, :\.1., & Hardingh:lln. T.E. (1996). ;\\;.Hrix protcoglycans, In W.O. Comper (Ed.) Ex/racdlulur .\lll/rix (Vol. 2.
pp. 200-229). Amsl(:rd~II11: Harwood ;\c;ldcmic Pubs.
Frank. E.H .. &: Grodzinsky. A.J. (198ia). Cartilage dcclromechanics-I. Ekclrokint:tic transduction nnd dfects of
pH .1lHI ionic strength. J Biol1lec!ulIlics, 30. 615_
Frank, E.H., & Grodzinsky, A.J. (1987b). Cartilnge electrollll'chnnics-ii. " continuulll model of l·;:u·tilagc declrokinetics and corn:l<:ltion with L'xperiments. J BiOlltt:chol1ics.
20, 629.
Freeman. M.A.R. (19i5). The fatiguc
c;lrtilage in the pathogenesis of ostcoanhro$is. :\cw OrtllOfJ Scal/d. 46, 313.
Fung, Y.C. (1981). Ouasi-Iinear \'iscch:lasticity oj" sol'ttisslles.
In Biolllcdlllllics: JleclulI1;cal Properlics of Lil'illg Tissues
(p, 226). New York: Springl'r·Veriag.
or
and ion:' lhrough ,\ POr<HIS-Pl'l"llll..·;lbll' ch;lrgl·d·h.nlrall
ti:-SliC. and slrl'~llnilig p(lll'llti;1! d'lla on nnrllwl h()\"inc ;1
licltlar Glnilagl..·. .Il3iol//<,clllll/;("S, 1r;. iU9.
Cll, "·,Y.. La;' W..\1. . .:i.: .'1 ow. V.c. (1997). ,\ triph:l"ic allaly:-.
or l1l'g,lli\'l: Olillloti<.· flo\\':" lhrough chMgL'd hydr,lll'd so
li:-SllCS. J BiOlIl,·c!Wllic.....W. i I.
CII. \\'.Y., Lti. \\'..\I.,..\: .\Io\\". \'.c. {199S).:\ llli\tllrL' thl·ory {
dl'trgl..'d h.nll';llCd sott lisSl!l..·s (ont;lil\ill~ lllultl-eIN
trolYll'S: PaSSI\'l' tr:lllspon and s\\'dling lll'h,,\·iors. J J3i
I1ICc!t EII~, 102. 169.
lubrkatlo11 for :t.\ial ;,.nnllll·lr'.\ Linder high !o;,ding cond
tiOIl~,.1 BioIlH·C!z(/uic.', 28. 1199.
Hodge. \\'...\ .. Fij:lll. R.S .. Clrlson,. 1\. .. 1.'1 al. (1%61. Conl.l
prl':-iSllrl' in till' human hip joinl Illl·~I:-illrl'd in yi\·o. Pro
Na/f :lclIe! Sci. US.--\. 83. 28i9.
Holmcs, ~1.H .. L.li, W.M .. l\: !\lo\\". V.C, (1985). Singular p~'
tllrb:ition ~Inalysis (Ill the nonlinear. f10\\".dqll'lldenl, COI
prL'ssi\'c slrl'ss-rl'laxalion bdl;wior of ~1I'lkul;lr (~lnihlgL'.
Biollll'ch I.:I/~, Wi. 206.
HOll. J.S., .\'Io\\". V.C .. L:ti. \\".\1 .. ('I al. (1992). :\n :In;d.\'sis
till' SqIH.'l'i'.L'-fillll lubrication llll'ch~llllsm for :1I·tleul:H· c~l
tilagl·. J Riollll'dul1Iic.... 15, 147.
Huhkr;llllZ. W. (I S9S). L'dwr dil' Spaltrichlllngl'll lkr G
lellkknorp..::J. \'~~l"lliIlldl/lll~ell dl'l' :llIrI/olI/iscln'lI C:;csdlsc!za
12. 1-+8.
~
I;~;::~~~{:;IA~. :"i &ill Curran.
P.F. (1975). NOI1t.'qllilibriIUJI ThaBiophyjics (4th l.'d.L Camhridgl': Han'anl
Prl.'ss.
G.£.. Tuke, M_A .• Dingle. J.T.. ct al. (1976). The dof proteolytic enZYllll's on thl.' Illl"chanical propt:.'l"ties
hunHlfl 'll·ticuIOlr canilagl.'. Biocltl!/IJ 13iopll,\'... :\cra,
I.
G.E. (1979). Mechanil.:i.ll propl.'rtics of <Inicular
In ~'I.A.R. Frl.'cman (E,L), .4t/lIlt :\aicillar C(lai·
cd .. pr. 333--t.l-1). Tunbridge Wells. U.K.: Pitman
v.J., Sah. R.I..., Grodzinsky. A.J .. ct al. (1994). :Vh:chani1'C!;ulalion of cartil;lge biosyllthetic b\:h~l\'ior: Ph~'sical
Arch Biocln'/It Biophy,,,', 311, 1.
Gu, W,Y.. & Mo\\', \!.e. (1998). On the conditional
<Qui"""",,·" of chemic:11 loading and ml'ch;lllical loading
articular C;lrt ilagc. 1 Biol/1('c:/wllics 3 I (111. 1181-11 $5.
W.:'\'I., & Mow, \I.e. (1978). Ultrafiltr,ltiol1 of synodal
'';'?:'"'''' by canilagc_ J Ellg .tlt:clt Dil· ..I8CI":.:. /0./. 79.
& Mo\\', r.e. (1980). Dnlg-induccd t'ompression of
,,''''ulal' cartil;\gt: during a permeation c.xpl,'rimenl. J Bio17. III.
W.M .. I-lOll, J.S .. &. Mo\\', V.C. (1991). A triphasic theory
the s\\'dling .lIld deformation beh~l\'iors of anicular
canil'a.". 1 Bicn/It'cll EII~. 1/3. 2-t.5.
& Saha. S. (1979). Cemenl line motion in bone. Sci-
204. 50 I.
,\'t., & \-Vt~iss, C. (1975). Rl'\'iew of nnicubr t·;\rtil;lgc
resl.'arch. Anhrilis Rhett/II, 18. 553.
(1968). Lubrication of animal joints: l. The mechJ BilJllli'Challics. ,. 193.
EC.. is: Radin. E.L. (1968). Lubric<llioll of :-lnimal joints;
III. The eHcc! of certain chtmic~\1 altnations of thl' cart iIag~ rind lubricant. .. \f/hi/is Rlh·UII/. J 1. 674.
F.C.. & Sokoloff. L. (1965). MO\·l.'IlH:nt and composition
of inlt:'rslitial fluid of cartilage. Arthritis Rhc.'ll!ll, S, 481.
H.. Ethcredgl', R.. &. Glimcher, M.J. (1975). In vitro
of articular c~ll"Iilage. I: I-Iydl'o'\~<prolinc, h(,'.'\and amino acid composition or bO\'inl: i.\rticular
c<l.nibge as It function of depth from thl.' surface: hydroxyproline cont"nl of the lubric;II11 and the weur debris as a
1ll~;lsure of wear. 1 B01l1! 10i1l1 Slfrg. 5i..t. 527.
Lipshiti'., H., Etheredge, R.. is: Glimcher, ~1.J. (1976). Changes
ill the he.'\os:lJlllne contt:nt and swelling ratio of articular
(:'Il·tibge as r11l1~lioll:; of depth from Ih" Sur!";,lC\..'. J BOlle
loilll Surg. 58..\. 11-t.9.
Lipshitl.. 1-1., &. Glimchcr. \1.J. (1979). In \'itro stud it'S of the
wenr of articular cartila!.!e. 1I'cm: 52, 297.
Malcolm. L.l. (1976) . ..t/l ~t'xpt:rill1ollal illl't':,tigatioll of 1111:
{i'jet/vllal al/{l d,,{or1lltlliOlUl! I'espOIIScs o( or/iclllar cartilage
illwrt,u:es 10 .wl/ic (lI/(/ dYl/(lIllic loadillg. Doctoral thesis.
Universitv of California. San Diego.
:\lankin. 1-1.:\., & Thrasher, A.Z. (1975). Water conlent and
bindin~ in normal and osteonrthritic human cartil:-lgc. J
BOlle
J~i/II Sun.:, 5i:\, i6.
Mansour. J.,\'I..& ~to\\", v.e. (19i6). The permeability or articu!;\r canilag..: ulU!l'r comprt'ssin:' str,lin lind at high pressures. J BOlle lohn SlIl',!;, 58;\, 509.
:\laroudas, A. (196611967). Hyaluronl( acid films. Proc IllS!
Jkch Ell!;. 1.011£1011,1811. 122:.
Maroudas, A. (1968). Physicochl.'llIical properties of c~\l·tilage
in Ii£ht of ion-exchangL' theory. Biophy,'; J. S. 575.
Maroud'IS. A. (1975.1. Biophysical chcmislr~' of cartilagin
ti~Slh:S with sllI.:cial refen::l1cC' 10 solute ~ll\d fluid transp
Biorhco!ogy. 12,233.
"taroud~ls. ..\. lI976). 8;1blll.:e bC(\\'l.'l.'1l swelling pn:ssun:
coll~lg"n tension in n(lrl1l~a1 alld ,k·gelll.'ratc c3nilage.
IIItc.',
260. 80$.
i\larou,!:Ls, A. (1979). Ph~'sicochcmkal propenies of anicu
cHrtilnge.ln M.A.R. Frl.'C'man (Ed.). Ildlill ,.\riiCl1ltu· Ca
la;:c: (2nd cd .. pro 215-290). Tunhridge Wdls. EnghllHI:
mall '\h:dical.
:\broud:\s, A.. \bchteL L, Grushkn, G.. \..'t al. (1991). Th\..
fecI of osmotil.: ~llld Ilh:ch;wical preSSllrl:S Oil Wi.ller pa
tioning in ~lJ'lit-lIl<lr (,:Hrtibgl·. Biocltt'l1I Biophys ..
10i3. 285.
\lcCutchcn, C.\\'. (19621. The frictional propi..·nies of :Ini
joints. it-'eM.5, 1.
,\kDevitt, C.A ..
Xluir. H. (1976). Biochemical chall~e
th,' cartilage of ill(' knce in eXpl.'rilll('nt~d <llld lIatul;d
h:o'lI·(hritis in Iht: do~. j BOIlt! 10ill/ S/lr~. 588. 9.J. .
.\-kachilll. G., & Ft:rgie. I~..\. (1975). Morph~logic~d pattern
articular cartilage fibrillation. J Pm/lal, 115, 231.
\'\(1\\', V.C .. AlllO~zky, S.P., &. Jadson, D.\V. {1992L K
Jlt'lIis('/ls: Basic ami Clillical F(J/(II(huiol/.~. N(,'\\, 'Y
Ra\'en Press.
,\10\\', V.e.. & Ateshian, G...\. (1997). Lnbrit~llion and \\"(,';1
dianhrodial joints. [n V.e. Mo\\' &. W.e. Ha~'t's (Eds.),
sic BiOUlt'c/wuics (2nd l'd .. pp. 175-315). Phibddph
Lippincot t-R~l\"l'n.
\lo\\', v.e" Atcshi'Hl, G.A .. Lai, W.M .. i..'l a1. (1998). Effl.'ct
fixed charges on the sll'ess-r(,'laXlltion bdl:l\"ior of h~'dra
soft lissues in a confint'd compression pl'Obkm. /I
Solid... & SIT/tc!/lres. 35, ~9.J.5-4961.
:\low, V.C., Gibbs. ,\I.e., LaL \\'.\1.. l.'t Ill. (1989a). Biphnsic
dentation of ~Il·ticular Glrlibge-Part II. A llurlh:ric:11 a
rithm ;\Ild an l'xpt.:riml.'U1;'ll stlld~'. 1 Bi()UI~'(:I/(l/1ic..;. 22, 8
'\'lo\\', V.C., Holmes, ;\1.1-1 .. &. LaL \\'.\'L (198"+). Fhlid tl'i.ll1s
and mechanical propl.'nies of articular c,lrtilage: A rev
J Hiol/lt'cha /lies. I?, 377.
\lo\\". \I.e.. Kuei, S.c.. LaL W.'\-1.. ('t al. (1980). Biphash: cr
and stress rdaxOIlion of anicul;ll' L·'lrtilage ill comprcss
Thl'ory and cxpl.'rimellts. J Biol1lt'dl EJlg. 102. 73.
Mow. V.c.. LaL W.;\I. . ..\: Rt.'dlcr. I. {197-H. Some surface c
aCIl.'ristics of articular cartilag~s. A SC;-\lliling electron
croscopy study and a th~'orl'\ic~d model ror the d.vna
inli,:ractioll of synovial fluid and articular cilrtilage. J
IIleclumics. 7. -t.-I9.
Mow, V.e. . ..\: R:llcliffi..'. :\. (1997). Structure :Hld funclion
articular c::Irtili1gc and menisclls. In v.e. ;"low l\! W
Hayi..'s (Eds.), B(I.'ii~· Ori"o!'(/I:dic: Biollll.·chal/ics (2nd
pp. lI3-11i). Philadelphia: LippincOIl-Ra\"l.'ll.
\!o\\', V.e., Zhll, W.B., Lai, W.M., ct al. (1989b). The in
ence of link protein stabilization on 1h..: viscoelastic p
ertici' of protcoglYC<1I1 aggregates. Bioclt':lll Biophys .992, 201.
:'I·luil". H. (1983). Proteogl~'cans as org:lnizl.'rs of the extra
lular matrix. Bim.:ht'JIl Soc TraJJs, I 1.613.
Myers. E.R .. Lai, \V.M .. &. ~·Iow. V.C. (1984). A continuulll
ory ;Illd an expcriment for the iOIl-inducl'd swelling beh
ior canila£c. 1 Biol1wdl r.II!.'" 106(21, 151-15S.
N\..'\\·b..:rry,
Zukosky, D.K:, & Hau!. R.C. (1997). Subf
tllrc insult to a knt:.·c joint causes alter~l(ions in the b
:-lnd in the rUflctional stiffness of o\'erlying cartilage. 1
Ihop Rc:s, 15. -t.50.
'*
W.N..
Onsagcr. L. (1931J. Reciproc:l1 relations in irn.'vcrsible
processL's. I. Phys ReI' 37, "+05.
Pau!. J.P. (1976). Force actions transmilted b~' joints in the
human bod\'. Proc Ror Soc Loud, /92B. 163.
Poole. A.R. (1986). Pr~)[coglycans in health and diseasc:
StrucLure and function. Hinchem J. 236. 1.
Radin. E.L.. and Paul, 1.1.. (1971), RespollSc of joints to impact loading. l. In vitro wear. Arthritis Rhellill. 14, 356.
Radin. E.L. (1976). Aetiology of osteoarthrosis. CIiIl RhculJI
Dis. 2, 509.
Ratcliffe. A.. &. ~'Iow, \f.e. (1996), Articular cartilage. In \V.D.
Compel' (Ed.), Extrace/llIlar Alatrix (Vol. 1, pp. 234~302).
AmsLcrdam: Harwood Academic Pubs.
Ratcliffc. A.. 'I\ler. J.. &: I-Iardingham. TE. (1986). Articuhlr
cartilage culture with inlerlcukin I: lncreascd releasc of
link pr~)tein. hyaluronate-binding region and other proteoglycan fragments. BiochclII J, 238, 571.
Redler, I.. &. Zimny. M.L. (1970). Scanning electron microscopy of normal and abnormal artlcuhll' curtilagc and
synovium. J BOl/c Joint SlIrg. 52/\. 1395.
Rcdlcr. L, Zimny. :'.'1.L., \Ianscl!, J .. et al. (1975). Thc ultrastructure and hiol11cchanical significancc of the tidemark
of articular canilagc. Clill Or/hop Rei Res, / /2.357.
Rosenberg. L.. Choi. H.U .. Tang. 1..-H .. et a!. (1985), Isolation
of dennatan sulf.ltc proteoglycans from mature bovine artindar cartilage. J Bio! Cht'lJl, 260, 6304.
Roscnberg, L.. Hellmann, \V., &. Kleinschmidt, A.K. (1975).
Electron microscopic studiL's or proteoglycan aggregatL'S
from bo\'ine articular cartilage. ./ Bio! Chew. 250, 1877.
Roth, V.. &: Mow. V.c. (1980). The intrinsic tensile bchavior of
the nl<\trix of bovine articular canilagc and its variation
with age . ./ BOHi.' ./oilu SlIrg, 62A, 1102.
Roughlev. P.J.. &. Whitc, R.J. (1980). Age-related changes in
t';e structure of thc proteoglycan sub~lnits from hum7'1ll articular cartilage. J BioI Choll. 255, 2!i.
Roughley. P.l .. \Vhite. R.l., & Santcr. V. (198 I). Comparison
of proteoglycans extracted from high- and low-weight
bearing human articular cartilage, with particular referencc to sialic acid content. J BioI Che/1/, 256, 12699.
Schinagl. R.M .. Gurskis. D., Chen, A.C., et a1. (!997). Depthdependent confined compression modulus of fullthickness bovine articular cartilage. J Ort/lOp Res. /5,499.
Schinagl. R.M., Ting. M.K .. Pricc. J.H .. el a1. (1996). Video
microscopy to quantitate the inhomogeneolls equilibrium
strain within articular cartilage during confined compression. illIll Biolllt'd Eng. 24. 500.
Schmidt, :\-;1.B., Mow, V.C .. ChUll. L.E., et al. (1990). Effects of
proteoglycan extraction on the tcnsile behavior of artiClIJar canilage . ./ Or/hop Rcs, 8. 353.
Schneiderman. R., Keret, D., &. Maroudas, A. (1986). Effects
of mechanical and osmotic pressure on thc rate of glycosaminoglycan synthesis in adult femoral head cartilage:
An in vitro study. J Orthop Res. 4, 393.
Schubert, i\'1.. &. Hamerman, D. (1968). A Primer 011 COIIlICCtive Tisslle Biochelllistry. Philadelphia: Lea &: Fcbiger.
Scott . .I.E.. & Orford, C.R. (1981). Dennatan sulphate-rich
protcoglycan associates with rat tail-tendon collagen at
the d band in the gap region. BiochelJl J, /97,113.
Selton, L.A., Gu, \V.Y., Lai, \V.M .. et a1. (1995), Predictions of
the swelling induced pre-stress in articular cartilage. In
A.P.S. Selvadurai (Ed.), .lleclulllics of Porous Atcdia (pp.
299-311). Kluwer Academic Pubs. Dordrecht. the Netherlands.
SCtlUIL L.A .. \leJ\\', V.e. \Iuller, F..J .. L'l :d. (199..+1. :\lcch<H
pnq)l:l'liL's of c:lninL' :lnictllar c;lrlil:tgL' :11"L' signil'kanLly
Il'rL'd following tr:lflSacliofl or IbL' ;lI'llL'l'ior l"l"tlCiak l
nh:nl. 1 Of/!zop Res. /2.451,
SL'llon. L.A., Toll.\':lIl1a. fl .. l\: \10\\'. V.e. (1998). S\\'L'lling
curling bl'ha\'ior oi aniL'ular canibge. J BiuJllt'c1/ r
/10.355.
Setton. L.A .. Zllli. \\'.13.. & :..,,10\\-. V.e. (19931. Thl' biph
poro\'iscol'l:lstic lk'h:l\'ior 01' anlcular cartib1:!c in C
prL'ssion: Role or thl' surracc zone. J Biollleclwl/ics.
581.
Sokoloff. 1.. (1063). ElasliciL,V of articular cartilagL': Ei'fcc
iOllS :lIld \'iscOlls solutions. Sci(l/ce. /4/. 1055.
Soltz, \1.A .. & A(l'shi:1I1. G.A. (1998). ExpnilllL'nlal \'nif
tion and Ihl'()]"L'lic,d prediction of canil:lgL' inkrsli
r1uid pressurizalion al an impl'rmL'able cont:lct intcr
in conl'inl'd corTlpn:ssion . .I Bio!!/cchanics. OCt 31 (
927-934
StoL:kWL'lI, R.S. (1979). Biolo.!!.y
COJ"/i!o!.!.c Cd/so Cambrid
C:\lllbridgl' Uni\'crsit.\' Prcss.
Sun, D.N .. Gu. \\'.1'.. Guo. X.E .. L·t al. (199S1. ThL' infhll'llcL
inhomogelleous fixed charge dL'l1sitv on canil,q!l' nlccha
L'leetrochclllical belw\'iors. '!i'mls On/lUp Res Soc. 23. 4
Swann. I).A .. Radin, E.L .. & Hcndn:lL R.B. (1979). Thl·lu
c:ltion of :lrliclIbr cartilage b~' syn(wial I'luid gl,\'co
tL'ins. Arthriris Rhein. 12, 665.
SW:\l1Il. fL\ .. Sih'L'r, F.H., Sla~'lL'r. H.S .. ct :d. {1%5L ThL' m
t'cubr structure ~lnd lubricating :lcti\'it,\' of lubricin f
bo\"inc and human s~'nO\'ial fluids. BiucltoJl J. 225. 19
S\\'cL'l. \l.B.E .. Thonar. E../.-\1.A., & \1:lrsh . .I. (1979). A
rL'lated ch,lngcs in protcoglYC:1I1 structurL'. Arch Hioch
or
Biopltys. / 98.
439~448.
Thompson, R.C .. Ocgcma. TR .. Lcwis . .1.1" L'l al. (1991 L
tcoanhrotic changcs al'tcr anHe Iransarticular load.
animal model. J BOHL' lohll Slirfi. 73:\. 990.
Thonar. E..I.-\L·\ .. Bjornsson, S., & KueHner. K.E. (19
Age.related changcs in canil:lge proteoglycans. In
Ku,-'ltner. R.S .. SclJle~'crbach. &. V. C. Hascall (Eds.). A
1111/1' Clirrilllg.c Biochcl!/islJY (pp. 273,-287). Nl'\\' Y
Ra\'en Prcss.
Torzilli. P.A .. &: \Iow. V.e. (19761. On the t'lInd:lI11elltal f
Iransport lllechanisms through normal and palhologic
tilage during function. 1. The formulation . ./ Biolllcch, 9
541-552.
Torzilli. P.A .. Rose. D.!:., & Dethl'mcrs, S.A. (I9821. Equ
rium watn partition in articular cartilage. Biorhw/ogy,
519.
Urban . .l.P.G .. & \-lc\luliin . .I.F. (1985). Swclling prcssur
thc intervertcbral disc: Influcnce of collagcn and pro
glycan content. Biorhc%gy, 22. 145.
Valhlllll. W.B .. StazmDc. E.J .. Bachrach. N.;vl . L'l a!. (19
LO<ld-controlkd compression of articular cartilagL'
duces a transicllt stimulation of aggn:can gene cxprcss
Arch Biocltc/!/ Biophys 353, 29.
Venn, \'I.F. (1978). Variation of chemical compOSition
age in hllman remora I head cartilagc. Ann RheulII Dis,
168.
\Vada. T. & Akiwki. S. (1987). An ultrastructural stud
solid matrix in articular cartilagc under uniaxial ten
stress. J .Ifill On//tJfJ Assn. 61.
\Valker. P.S .. Dowson. D.. Longfeild. \I.D .. et al. (19
"Boosted lubrication" in syno\'ial joints b~' rluid ent
lllent and enrichment. :Inn !VIC/IIII Dis, 27, 511.
Walka. P.S .. Unswonh, :\ .. Dowson, D.. I,:t a!. (1970). ~-lod(.' of
ag!!rcgation of hyaluronic acid prntdn cOll1pk:x on tilL'
sll~facc of articular cartilage. .. \/11I Wlt'lflll Dis. 29. 591.
\V;JOg. C.B., &. Mow, V.C. (199S). Inhomogendty of aggreg,HL'
n;odulus affects c~rtil:tgL' cornpn.:ssi\·I..' strcss-n:la:Xi1tion
behavior. 1i'alls Or/hop Uc:s Soc, 13( I). 481.
Weiss.
Roscnbl.:rg. L., & J-1e1fL'1. A.J. (1968). Alluhraslrllctural study of !lumlal young adllli hum;11l anicular cani·
lage. J BO//(: 10illl SlIrg, 50A, 663.
\Villiams, P.F., j1owdl. G.l.., &: Laberge, M. (1993). Sliding
friction analysis of phosphatidylcholinl: as a boundary
lubricant for anicul<"lr cartilage. Prot Ills 1 l/ech EIl.!jI'S,
c..
l07, 59.
Woo. S.L.-Y., Mow, V.c., &: Lai, W.~,t. (1987). l3iorncchanical
propcnics of ;Irticubr cartilage. In /-IllIu/book of Bioengineering (pp. 4.1-.....·H. New York: McGrmv-Hill.
Woo, S.L.Y.. LL'wsay G.A., Runco, 1.1., ct:ll. (199;). Struc
aml function of (t:ndolls and ligalllt:nts. In V.C. ~·1o
W.e. Hayes (Eds.), Basic Orthopaedic BiollleclUlllics (
cd., pp. 109-2.5 I). Phibddphia: Lippincolt-Raven.
Zhu, \V.B .. Iatridis, J.e.. Hlibczjk. V.. Cl al. (1996). DCle
n;ltioll of collagen-prot('oglycan interactions in vitr
BioJl1cclwnics. 29, 7i3.
Zhu, \-V.B., Lli, \V.~-I., &: J'I·low. v.e. (1986). Intrinsic qu
linl..'ar viscoebstic bL'havior of thc c.\tracelllllut' m;:Hrl
cartilagL'. 7i'tIIIS OrthofJ Res Soc, 1/, 407.
Zllu, \V.B., Lai. W.i\1., &. I\\OW, V.e. (1991). TIl(: dl.'nsity
sll"L'ngth of proteoglycan-proli..'oglycan interaction site
C(IIlCenlrakd solutions. J Biolllt'C!liIllics, ],1, 1007.
Zhll, W.B., ,\In\\', V.C .. Koob. '1'.1 .. ct al. (1993). Viscoela
shl.."u· propi..'rtil..'s of articular cartilagL' and the <:ffcc!:.
glycosidase treatments. 1 Ortlwp Res, I J. 771.
AIRTI'CULARCARTILAGE
,streSses sustlllned by cont::ict joint surfaces
." '.' '"
".",.'.::~::~><,,:,>~'
To provldoa smoothwear·reSistantbearing
s'urfac~ecrease b1ction~'
'~:,',
'.•···HI.,hlvSpeclalizedTlssu'e
~
;'
I
'I
EXTRACELLULAR COMPONENTS
CELlULAR COMPONENTS
Extracellular M3trix-ECM-
1
\~
,-'/,.,'-'
"
POROUS PERMEABLE SOUD PHASE
.!
-Qrganlt Component--:
i
.A
;~TERS~IT.I~]f3~!·[)rHASE ,:;;;
~!ceir'~~~~!%!~i;'1f'~~~&!,i~'>
0,
:":1
~
j
:1
Biph;uic Str'uc(urc
II
:,:,'1
BIOMECHANICAL PROPERTIES
\!
Animtropy,Viscocl;micity. Swelling Bchavior-ColllprCHivc lo~d bearing c~pJcity-
.~,
FLOW, CHART 3·1
Articular cartilage structure and biomechanicaJ properties. ~
'~\j
"Thh How ChMl is
d~si9ncd
fOI das>room or group
disCLl~~io".
Flow than
i~
not mCi'lr'lt \0 be cxh<lustive.
Results in
Disruption of collagen-PG solid matrix
PG "wash out"
Gross alteration of the normal load cartilage mechanisms
~fLOW CHART 3'2", Articular cartilage wear mechanisms. ~
"This flow chart is designed for c1assroorn or group discussion. Flow ch.Jr1 is not meant to be exhaustive.
~
;;;
E
~
-"!
.2
-.;
v
~x
~
~'
\J
'"C'
~
v
>-
'"0
0
~
Ci
ei
0
>
't;;
.
0
0
0
0
.0
8
~
E
5
E
0
0
.;0
~
.-
~
.,;
c
'"
~
."
~
'"
~
.;0
~
v
£
.~
."
~
~
·v
~
~
~
E
V
."
~
e:;;
."
0
,
0
~
's1;' <;-S
:E
:;
,
!1.
g
.~
~
'0
~
~
"'6
E
0
e
•
0
v
~
"
~
iii §
0
."
g
~
0
C
.~
0
."-
,J.'
t
,
0
.,;
~
i=
•
!
.';".'i~'-
.;.
!j
'.:
./
!
Biomechanics of
Tendons and
ligaments
Margareta Nordin, Tobias Lorenz, Marco Campello
Introduction
Composition and Structure of Tendons and Ligaments
Collagen
Elastin
Ground Substance
Vascularization
Outer Structure and Insertion Into Bone
Mechanical Behavior of Tendons and Ligaments
Biomechanical Properties
Physiological Loading of Tendons and Ligaments
Viscoelastic Behavior (Rate Dependency) in Tendons and ligaments
Ligament Failure and Tendon Injury Mechanisms
Factors That Affect the Biomechanical Properties of Tendons and
Ligaments
Maturation and Aging
Pregnancy and the Postpartum Period
iVlobilization and Immobilization
Diabetes lvIellltus
Steroids
Nonsteroidal Anti-Inflammatory Drugs
Hemodialysis
Grafts
Summary
References
Flow Charts
•
Introduction
The three principal structures that c10selv surround,
connect and stabilize the joints of the ~keletal svst~Il1are tendons, ligaments, and joint capsules. ~\l­
tIlollgh these structures arc passive (i.e., thev do not
acth'ely produce motion as do the l11usck:s), each
pl.iJ's an essential role in joint motion.
The role of the ligaments and joint capsules,
\vI11c.h connect bone with bone, is to auament the
I-IICchanical stability of the joints, to gl1i(iL~joint motion, and to prevent excessive motion. Ligaments
anel joint capsules act as static restraints. Tile functicm of the tendons is to attach muscle to bone and
to transmit tensile loads from muscle to bone,
thereby.' producing joint motion, or to maintain the
hpcly posture. The tendons and the muscles form
the muscle-tendon unit, which acts as a dvnamic n>
straint. The tendon also enables the mtls~le bell\' to
p¢at an optimal distance from the joint on whi~h it
acts without requiring an extended length of muscle
~
between origin and insertion.
Tendon and ligament injuries and derangements
a~'e common. Proper management of these disor(leI'S requires an understanding of the mechanical
properties and function of tendons and ligaments
ZtIld their capacity' for selF-repair. This chapter discusses the following:
Composition and structure of tendons and
ligaments
Biomechanical properties and behavior of
normal tendon and ligament tissue
Biomechanical properties and behavior of injured tendon and ligament tissue
tutes a large portion of the organic matrix of b
and ~artilage and has a unique mechanical supp
ive function in other connective tissues such
blood vessels, heart, ureters, kidneys, skin, and li
The great mechanical stability of collagen gives
tendons and ligaments their characteristic stren
and flexibility.
Like other connective tissues, tendons and l
ments consist of relatively few cells (fibroblasts)
an abundant extracellular matrix. In general,
cellular material occupies approximatelv 2000 of
total tissue volume, while the extraccliular ma
accounts for the remaining 80°/(;. Approxima
70t;0 of the matrix consists of water, and appro
mately 30% is solids. These solids are collag
ground substance, and a small amount of c1a;
The collagen content is generallv over 75 % and
somewhat greater in tendons than in ligame
(Kassel', 1996); in extremitv tendons, the solid m
rial may consist almost e~1tirely of collagen (up
99% of the dry weight) (Table 4- I).
The structure and chemical composition or l
ments and tendons are identical in humans and
many animal species such as rats, rabbits, dogs,
monkey's. Hence, extrapolations regarding~ th
structures in humans can be made from the res
of studies on these animal species.
COLLAGEN
The collagen molecule is synthesized hv the fl
blast \vithin the cell as a la;'ger precurso·r (proco
gen), which is then secreted and cleaved extrace
larl y' to. become collagen (Fitton-Jackson, 19
(Fig. 4-1). Tendons and ligaments, like bone,
composed of the most common collagen molec
Several factors that affect the bioI11echanical
u t <:>
function of tendons and liuaments
are auin
co
"" co' ')fe'(1nancy, mobilization and immobilization, diabetes,
nonsteroidal anti-intlammatory drug (NSAID) use,
and the effects of hemodialvsis. Biomechanical considerations regarding graft~ are also given.
Tendon
I Cellular Material:
Composition and Structure
of Tendons and Ligaments
Tendons and ligaments are dense connective tissues
knc)\vn as parallel-flbered collagenous tissues. These
sl:arsely vascularized tissues are c0l11posed largelv
of- collagen, a fibrous protein constituting approximately one third of the total protein in the bodv
(White, Handlel~ & Smith, 1964). Collagen consti-
i
I EX~~:~~I~'ar Matnx
I
20%
20%
Fibroblast
80%
80%
60-80%
60-80%
Solids:
20-40%
20-AO%
Collagen:
70-80%
slightly hig
Type 1
90%
9S-99%
Type 3
10%
1-5%
Ground substance
20-30%
slightly
En(l0!0f!On
f:1l1nd!e
ProcoHagen
Fibroblasl
Collagen
molecules
- - - - - - - - - - - - - _ . _..._._----
-
_ _......
----._.--.~---
._--._---------
Schematic representation of wllagen fibrils. fibers. and bun-
f'xtrac('lIuiar lll<1!rl)( ill cl PiHt111e-! ilflangelll('llt {o form m;·
dles in tendons and collagenous ligaments (not drawn to
crofibrils clnd then fibrils, The staggered array of the mole
(tIles, in which each overlaps the other, gives .1 banded ap-
scale). Collagen molecules. triple helices of coiled polypep-
tide chains. are synthesized and secreted by the fibroblasts.
These molecules (depicted with "heads" and "tails" to repre-
sent positive and negative polar charges) aggregate in the
type I collagen. This Il101eculc consists of three
polypeptide chains (ex chains), each coiled in a lefthanded helix wilh approximately 100 amino acids,
which give it a total molecular weight or approxi.
mately 340,000 daltons (Rich & Crick, 1961) (Fig.
4-2). Two or the peptide chains (called ,,·1 chains)
are identical. and one differs slightly (the 0:-2 chain).
The three a-chains are c0l11binccl in a right-handed
triple helix. which gives the collagen molecule a rodlike shape. The length or the molecule is approxi.
mately 280 nanOllleters (nIT}), and its diameter is approximately 1.5 nm.
Almost two thirds of the collagen molecule consists of three amino acids: glycine (33%), proline
(15%), and hydroxyproline (15°/(;) (Ramachandran,
1963). Every third amino acid in each 0: chain is
glycine, and this repetitive sequence is essential for
the proper rormation of the triple helix. The small
size 01' this amino acid allows the tight helical pack·
ing of the collagen molecule. C'vlorco\'cr. glycine enhances the stability of the molecule by forming hydrogen bonds among the three chains of the
superhelix. Hydroxyproline and proline form hydrogen bonds, or hydrogen-bonded wal~r bridges.
within each chain, The intra- and inlcrchain bonding. or cross-linking. between specific groups on the
chains is essential (0 the stability of the molecule,
Cross-links are also formed between collagen
molecules and are essential to aggregation at the
,0
the Coll<1gen iibri!s under the electron
pcarancE'
llIicro,:>wpe. Tilt:' hb!!l ... <)~J9rcgate further into hbers, ~"... I1l<:h
(omc together into dcn<,cly piicked bundles.
lihril k'\'l'l. IL i...; till.' I,.'I'{):-;:,,-!illkl,d dl~II·~h.:[l'I' ur Llll'l'
1~lgl'll lihrib ikll I~h'l'~ ,"'II'l'll~lh [0 thl' li:--:-il.l('.'" IIl
C( lmpOSl' and <:t1I< I\\':-- I hl.':--l' I b:"Ul'S I( I l"tl1"ll.·1 il III und
1ll1.'("II<tllic~d :--lrl':":--. \\'illlill 111(' lihril:-., !Ill.' mulecul
<:tl"t' ;Jppar('Il!I~' tTUss-lillked h.y "hl'<:\(.I·I(l-i~lil" illte
<:ll:tiulb (Fi~_ -l~ I L hUI inlcrlihrilbr IT(I:-..... -linkill~ o
rnOl"L' complL'\ n~llllr(' <:d~(l rn~l.'" on·lll'.
III 111..'\\-1\- fcwlllt:d ,,:(lll~l~l'l1. thl' CI'(l:-.:--·lillk:-- ~\1\':' r
:lli\"\,.,h- k\\' :lnd ~lrt' rl.~dlll.'ihl\:: 1111.' l.'(llb~l.'l1 b :,>olll
ill rJt'lIlral sail :"ohlliol"l:' <:!llll ill ~Icid :"(IIUliull:--, a
1Ilt' i.:ro:--~-Ii111..:-; a I"l' hI i rI., t'a~i I.,' dL'Il~llll rt'd hy hl.'at.
t.:olla~t'n a~L':-', lilt' 1111<:11 rlullIhtT nl rl'lilldhk' tTOS
link:-. dl'LTl'a:-'L':-' lei ~l tllinilllurn ~b ~l 1~lr~l' lllllllhl'r
:--(~lhk', nonrl,dul'ihk· lTCls:"-lird,:,, ~lrl' fOI"t'lll'd. \bIUI
l'<"lI~l~t'n i~ nul sU[lIhk, ill Ilt'lllral :"i.dl ~(lIt1ti(Jll:" 01'
~ldd solulio1\:",
and il
:-illlyi\'l'~:\ hi~llL'l' dl'I1;,lllll'~l
tl'l1IjWl':lllll'l', (1"-"0[' ~\ I\~\'il'\\'
or
lTo:"s-link:\1!l' ill ~'()
12l'll, Sl.'l' Viidik, D~lllil'lsl'll, t~ Chlulld, 10~2,l
.:\ lihril is !"(Irllll,d h," till' a~~I·q.!.ali(l1l
Sl'\'t'l
l,'olla~l'1l lllolL'ctlk's ill ;1 qU<lltTn~lIY Slrtlc!url.'. Th
strllclllrl.', in \\·Ilh.:h t"ach Illnkndt' O\'t'!·bp:-. I
otht'r, is I\'spollsihk for 1I1l' J'L'lk'alillt! IXIII((:-; o
:':'t'J"Yl,d Oil thl' lihril", lllH..kT till' t,k-l'II'OIl lllilTOS(O
or
(Fi~,
--l-2: ~l'l' :ds(l Fig, "'''''). Till' qll,llt'l'll:II-," Slrll...: lu
or l.T)lIa~cn
rdalt'S 10 till.' or~~1I1i:t<:lli(l1l III l.·ollagl.:
lllolt'ClIll's il110 :l :"l~lbk. 10\\' l.'lll'r~i,:lic bi(llo!!ic
unit ha:--t'd 011 :J rq:!.ul<:\1· a:':"(ll"ial iOll (~r ~\(,.l.iaCt'lll ';11
t'''.'lIks' h~ISil' <:llId ~lcidil' :llIlino ~ll·ids. B., :In:ll1gi
~\d.iacl'lll col1at!.I.'l't 1I1pk'l:uks in a qLl~lnl'l'-st~lg1:-~t'r. o
om
64
Fibril
000
, 0
or;:
Overlap zone
010
r
;:r:
Packing of
'molecules
I
I
Hole zone
Micrafibrils
~~~~=~'~'~'~~~~~~~~~
I
I
.................
..-",/
~~/
/
...
-'
...........
rCollagen r'.'='='=====:2~80~nm~======):J~
i' molecule
!...
I
........ .J
. ,.
,
-
,
!
..................
"
.......................
.......... i
a.2~~~t
1.5 nm
~,
1
Schematic drawing of collagen microstructure. The collagen molecule consists of three alpha chains in a triple helix
(bonom). Several collagen molecules are aggregated into
a staggered parallel array. This staggering, which creates
hole zones and overlap zones, causes the cross-striation
"
(banding pattern) visible in the collagen fibril under the
electron microscope. Adapled from Prockop, 0.1.. & Guzman,
;\! A. (1977). Col/agen dise,)ses and the biosynthesis of collagen
Hasp Pract. Dec. 61-68.
positely charged amino acids are aligned. This stabl.c structure will require a great amount of energy
,':~ri,d force to separate its molecules, thus contribut:.ipg to the strength of the structurc. In this way, orgU)lizcd collagen molecules (five) form units of microfibrils, subfib"ils, and fibrils (Fig. 4·3) (Simon,
1994). The fibrils aggregate f'urther to form collagen
fibers, which are visible under the light microscope.
;;./(~hese fibers, which range from I to 20 J.L111 in cliam''.;:,'',¢ter, do not branch and l11av be many centimeters
"j<lI1g. They reflect a 64-nm periodicity of the fibrils
.:!:
and have a characteristic undulated form. The fibers
aggregate further into bundles. Fibroblasts are
aligned in rows between these bundles and arc elongated along an axis in the direction of ligament or
;'i':. ~.enclon function (Fig, 4-4).
';/: The arrangement of the collagen j-ibers differs
somewhat in the tendons and ligarncnls and is
Suited to the function of each stru~ture. The fibers
composing the tcndons have an orderl.y, parallel
arrangcn)cnt, which cquips the tendons to handl
lhe high unidirectional (uni4lxial) lensile loads t
which they are subjected during activity (Fig. 4-4A)
The Iigamenls gencrally sLisrain tensile loads in on
predominant direction but may also bear smalle
(cnsile loads in other directions; their fibers may no
be completely parallel but arc closely interiaccd
with one another (Fig. 4-48). The specific orienta
tion of the fiber bundles varies to some cxten
among the ligaments and is dcpendcnl on the func
tion of the ligament (Arnie! et al .. 1984),
The metabolic tUI'nover of collagen may be stud
ied by tritium labeling of h)-!droxyprolinc or glycin
and by autoradiographic methods, Studics in ani
mals have shown that the half-life of collagen in ma
lure animals is very long: the same collagen mole
cules may exist throughout the animal's adult lire
however, in young animals and in physically altered
(e.g., injured or immobilized) tissue, the turnover i
acceleratcd. Rabbit studies have shown melaboli
activity to be somewhat greater in ligaments than in
tendons, probably because of different stress pat
terns (Amie! et al .. 1984).
ELASTIN
The mechanical properties of tendons and liga
ments are dependent not only on the architectur
and properties of the collagen fibers but also on th
proportion of elastin that these su'uctures contain
The protein elastin is scarcely present in tendon
and extrcrnily ligaments, but in elastic ligamcnt
such as the Iigamenlum Oavum, the proportion o
elastic fibers is substantial. Nachemson and Evan
(1968) found a 2 to I ratio of clastic to collage
fibers in the ligaments flava. These ligamcnls, which
connect the laminae of adjacent vertebrae, appea
to have a specialized role, which is to protect th
spinal nerve roots from mechanical impingement
to pre·stress (preload) the motion segment (th
functional unit of the spine), and to provide som
intrinsic stability to the spine.
GROUND SUBSTANCE
The ground substance in ligaments and tendons con
sists of proteoglycans (PGs) (ur> to approximatel
20% of the solids) along with stnlctlll,JI glycoprotcins
plasma proteins, and a variety of srnall n1olcculcs
The PG units, macromolecules composed of variou
sulfated polysaccharidc chains (glycos<lminoglycans
bonded to a corc protein, -bind to a long hyaluroni
acid (HA) chain La form an cxtremely high molecula
Fibroblasts
Crimp
Fascicular membrane
Schematic representation of the microarchitecture of a tendon .
•
•
Nearly parallel
bundles of
collagen fibers
Parallel bundles of
collagen fibers
I~=
~
Fibroblasts
ligament
Tendon
A
B
Schematic diagram of the structural orientation of the
fibers of tendon (A) and ligament (8); insets show longitu-
dinal sections. In both structures the fibroblasts are elongated along an axis in the direction of function. Adapted
from Snell, R.5. (984), Cfinical and F~ln(tional Histology for
Medical Stuejents. Boston: Liule. Brown
weight PC aggregate like Lhm found in Lhc ground
substance of anicular cartilage (set.:' Fig. 3-6).
The PC aggregates bind most of the extracellular
water of the ligament and tendon. making the matrix a highl~' structured gel-like material rather than
an amorphous solution. Furthermore. by acting as a
ccment-likc substance betwcen the coHagen microflbrils. they ma),' help stabilize the collagenous
skeleton of tendons and ligaments and contribute to
the o\'crall strength of these composite structures.
Only' a small number of these molecules exist in tendons. however, and their imporlance for its biornechanical properties has been questioned,
VASCULARIZATION
Tendons and ligaments have a Iimitcd vascularization. which alTeets dircctl~' their healing process [lnd
metabolic activit~;. Tendons receive their blood supply dircctly from \'cssels in the perim.vsiulll. the
periosteal insertion. and the surrounding tissuc via
vessels in the para tenon or mesotenon. Tendons surrounded b~) paratenon have been referred to as vascular tendons. and those sllrround~d by a tendon
sheath as avascular tendons, In tendons surrounded
by' a paratcnon, vessels enter rrom many points on
the pcripher.v and anastomose with a longitudinal
svstem or capilbries (Fig, 4-5),
India ink-injected (Spalteholz technique) into the calcaneal
tendon of a rabbit, illustrating the vasculature of a
paratenon-covered tendon. Vessels enter from many points
on the periphery and anastomose with a longitudinal system
of capillaries. Reprinted wirh permission from Woo, S.L. Y, An,
K.N., Amoczky, D.V.M., et at. (1994). And!Om}~ biology. and biomechanics of rhe iendon, ligament, and meniscus. In S.R. Simon
(Ed) Orthopaedic Basic Science (p. 52). Rosemont, Ii: MOS.
··f
"
•
The vascular pattern for tendons surrounded by
tendon sheath is different. Here the meso tenons a
reduced to vincula (Fig. 4-6). This avascuJar regio
led a variety of researchers to propose a dual pat
way for tendon nutrition: a vascular pathway, an
for the avascular regions, a synovial (diffusio
pathway. The concept of diffusional nutrition is
primary clinical significance in that it implies th
tendon healing and repair can occur in the absen
of adhesions (i.e" a blood supply). Conversely, lig
ments in comparison with surrounding tissue a
pear to be hypovascular. However, histological stu
ies reveal that throughollt the ligament substan
there is a uniform multivascularit)', which orig
nates from the insertion sites of the ligament. D
spite the small size and limited blood flow of th
vascular s)'stem, it is of prima(y importance in t
maintenance of the ligament. Specifically, by pr
viding nutrition for the cellular population. this va
cular system maintains the continued process
matrix synthesis and repair. In its absence, damag
fTom normal activities accumulates (fatigue) an
the ligament is at risk for rupture (Woo et aI., 1994
Ligaments and tendons have been shown in bo
human and animal studies to have a variety of sp
cialized nervc endings and mcchanorcceptors, Th
play an important role in proprioception and noc
ception, which are directly related to the functio
ality of joints.
the tendon substance. ReprinlGd wirh permission from Woo,
S.L. Y, An, K.N., Amoczk}~ D. VM., et al. (1994). Anatomy, biolog
and biomechanics of (he rendon, ligament, and meniscus, In S.R
Simon (Ed.). Orthopaedic Basic Science (p. 52). Rosemont. IL:
MOS.
OUTER STRUCTURE AND INSERTION
INTO BONE
Certain similarities arc found in the outer structure
of tendons and ligaments, but there are also important differences related to function. Both tendons
and ligaments arc surrounded by' a loose areolar
connective tissue. In ligaments, this tissue has no
specific name, but in tendons it is referred to as the
para tenon. More structured than the connective tissue surrounding the ligaments, the paratcnon forms
a sheath that protects the tendon and enhances gliding. In some tendons, such as the flexor tendons of
the digits, the sheath runs the length of the tendons,
and in others the sheath is found only! at the point
where the tendon bends in concert \vith a joint.
In locations where the tendons are subjected to
particularly high friction forces (e.g., in the palm, in
the digits, and at the level of the wrist joint), a parietal synovial layer is found just beneath the
paratenon; this synoviun1-like membrane, called the
epitenon, surrounds several flber bundles. The synovial fluid produced by the synovial cells of the
epitenon facilitates gliding of the tendon. In locations where tendons are subjected to lower friction
forces, they are surrounded by the panltenon only'.
Each fiber bundle is bound together by the endotenon (Fig. 4-1), \vhich continues at the musculotendinous junction into the perimysium. At the
tendo-osseous junction, the collagen fibers of the
endotenon continue into the bone as Sharpey's perfOl·ating fibers and become continuous with the
periosteum (Woo et aI., 1988).
The structure of the insertions into bone is similar in ligaments and tendons and consists of four
zones; Figure 4-7 illustrates these zones in a tendon.
At the end of the tendon (zone 1), the collagen fibers
intermesh with fibrocartilage (zone 2). This fibrocartilage gradually becon1es Inineralized fibrocartilage (zone 3) and then merges into cortical bone
(zone 4). The change from more tendinous to more
bony n1aterial produces a gradual alteration in the
mechanical properties of the tissue (i.e., increased
stiffness), which results in a decreased stress concentration effect at the insertion of the tendon into
the stiffer bone (Cooper & Misol, 1970).
Merhanical Behavior of
Tendons and Ligaments
Tendons and ligaments arc viscoelastic structures
\vith unique mechanical properties. Tendons are
Electron micrograph of a patellar tendon insertion from
dog, showing four zones (:<25,000). Zone 1, parallel coll
gen fibers; zone 2, unmineralized fibrocartilage; zone 3,
mineralized fibrocartilage; zone 4, cortical bone. The lig
ment-bone junction (not pictured) has a similar appearance. Reprinted with permission from Cooper, R.R. & Misol, S
(1970). Tendon and ligament insertion. A light and electron
microscopic study J Bone Joint Surg, 52A, I.
•
strong enough to sustain the high tensile forces th
result from muscle contraction during joint moti
yet are sufficiently' flexible to angulate around bo
surfaces and to deflect beneath retinacula to chan
the flnal direction of muscle pull. The ligaments
pliant and Oexible, allowing natural movements
the bones to which they attach, but are strong a
inextensible so as to ofTel' suitable resistance to
plied forces.
Analysis of the mechanical behavior of tendo
and ligaments provides important information
the understanding of injury mechanisms. Bo
structures sustain chiefly tensile loads during n
mal and excessive loading. When loading leads
injury, the degree of damage is affected by' the r
of impact as well as the amount of load.
VIVI~,-."MI."."'L PROPERTIES
means of analyzing the biomechanical properof tendons and ligaments is to subject specimens
tensile deformation lIsing a constant rate of elonThe tissue is elongated until it ruptures, and
resulting force, or load (P), is plolted. The resultload-elongation Clll-ve has several regions that
the behavior of the tissue (Fig. 4-8).
The first region of the load-elongation cuniC is
the "toe" region. The elongation reflected in
region is believed to be the result of a change
the wavy pattern of the relaxed collagen fibers. In
region, the tissue stretches easily, without much
and the collagen fibers become straight and
their \vavy appearance as the loading proa",:sSl:s(Hirscb, 1974; Woo et a!., 1994) (Fig. 4-9, A
B). Some data suggest, however, that this elongamay' be caused mainly by' interfibrillar sliding
4
3
2
I
1mmLI
I
I
I
Elongation (%»)
_
Load-elongation curve for rabbit tendon tested to failure
in tension. The numbers indicate the four characteristic regions of the curve. (1) Primary or "toe" region, in which
the tissue elongated with a small increase in load as the
wavy collagen fibers straightened out. (2) Secondary or
"linear" region, in which the fibers straightened out and
the stiffness of the specimen increased rapidly. Deformation of the tissue began and had a more or less linear relationship with load. (3) End of secondary region. The load
value at this point is designated as PI;,,' Progressive failure
of the collagen fibers took place after PI'" was reached, and
small force reductions (dips) occurred in the curve. (4) Maximum load (Pm.,,) reflecting the ultimate tensile strength of
the tissue. Complete failure occurred rapidly, and the specimen lost its ability to support loads. Adapted from Carls/edt,
c.A. (7987). Mechanical and chemical factors in tendon healing:
Effects of indomethacin ancl surgery in the rabbit. Acta Orthop
Scand Suppl, 224.
~~---------
Scanning electron micrographs of unloaded (relaxed) a
loaded collagen fibers of human knee ligaments
(x 10,000). A, The unloaded collagen fibers have a wavy
configuration. S, The collagen fibers have straightened
under load. Reprinted with permission from Kennedy, J. C,
Havjkins, R.I. Willis, R.B., et al. (7976). Tension stuclies of hu
man knee ligaments, Yield point, ultimate failure, and disrup
of the erueiare and tibial collateral ligaments. J Bone Joint Su
S8A.350.
and shear of the interfibrillar gel (ground substan
(for review, see Viidik, Danielson, & Oklund, 198
As loading continues, the stiffness of the tis
increases and progressively greater force is requi
to produce equivalent amounts of elongation. T
elongation is often expressed as strain (e), whic
the deformation of the tissue calculated as a p
centage of the original length of the specimen
strains arc increased (strain values of between 1.5
and 4% [Viidik, 1973]), a linear region will follow
the toe region. This sudden increase in slope represents the second region in the diagram and con"csponds to the response of Ihe tissue to further elongation (Diamant et al.. 1972).
Following the linear region. at large strains the
stress-strain ClIlve can end abruptly or curve downward as a result of irreversible changes (failure)
(Woo et aI., 1994). Where the curve levels off toward
the strain axis, the load value is designated as Plin ,
The point at which this value is reached is the yield
point [or the tissue. The energy uptake to PHn is represented by the area under the ClIl·ve lip [0 the end
of the linear region.
\Vhen the linear region is surpassed, major failure of fiber bundles occurs in an unpredictable
l11anne1'. \Vith the attainment of maximum load that
reflects the uhimatc tensile strength of the specimen, complete Failure occurs rapidly, and the loadsupporting ability of the tendon or ligament is substantially reduced.
The modulus of elasticity for tendons and liga~
ments has been determined in several investigations
(Fung, 1967, 1972; Viidik, 1968). This parameter is
based all a linear relationship between load and deformation (elongation), or stress and strain; that is,
the stress (force per unit area) is proportional to the
strain:
elastic deformation that they endure under tensi
strain and the storage and loss of energy. Durin
loading and unloading of a ligament between tw
limits of elongation, the elastic fibers allow the m
tcrial to return to its original shape and size aft
being deformed. Meanwhile, part of the enelspent is stOt"cd; what is left will represent lhe energ
loss during Ihe cycle and is called hysteresis. Th
area enclosed by the loop represents the energy lo
(Fig. 4-1 I).
PHYSIOLOGICAL LOADING OF TENDONS
AND LIGAMENTS
The ultimate tensile strength (P,,,:..,,) of ligamen
and tendons is of limited interest from a function
standpoint because under normal physiologic
conditions in vivo these structures are subjected
a stress magnitude that is only approxirnately on
third of this value. The upper limit for physiologic
strain in tendons and ligaments (when running an
jumping, for example) is from 2 to 5% (Fung, 1981
E = rIlE
where E = modulus of elasticity
(T
=
E =
stress
strain
In the toe portion of the load-elongation curve (or
the modulus of elasticity is not
constant but increases gradually. The modulus stabilizes in the fairly linear secondary region of the
curve.
The load-elongation CUPie depicted in Figure 4-8
generally applies to tendons and extremity ligaments. The curve for the ligamentum flavulll, with
its high proportion of elastic fibers, is entirely different (Fig. 4-10). In tensile testing or a human
ligamentum flavlIIll, elongation of the specimen
reached 50% before the stiffness increased appreciably. Beyond this point. the stiffness increased
greatly with additional loading and the ligament
failed abruptlv (!'eached P"",,), with little further deformation (Nachemson & Evans, 1968).
The proportion of elastic proteins in ligaments
and capsules is extremely important for lhe small
stress~slrain curve),
~1.'·
o
-70
Elongation
(~o)
Load·elongation curve for a human ligamentum flavum
(60 to 70% elastic fibers) tested in tension to failure. At
70% elongation the ligament exhibited a great increase
stiffness with additional loading and failed abruptly with
out further deformation. Adapted from Nachemson, A.L., &
Evans, J.H. (1968). Some mechanical properties of tile third hu
man lumbar imerlaminar ligamenr (ligamentum flavum). ) B;omech, 1, 21/-220.
•
,--? .
1 cycle
~Peakload
~
Elongation
Typical loading (top) and unloading curves (bottom) from
tensile testing of knee ligaments. The two nonlinear curves
form a hysteresis loop. The area between the curves, called
the area of hysteresis, represents the energy losses within
the tissue.
'0 ,Few studies or loading of tendons or Ii!!aments in
\:ivo have been perrorm~d. Kear and Smith (1975),
using the strain gauge method, measured the rnaxinUll strain in the latcral digital extensor tendons of
sheep. The strain reached 2_6t}~ while the shl:cp
were trolling rapidly and decreased when the trolting speed decreased. This maximal strain occurred
for only 0.1 second during each stride. The maximal
load imposed on the entire tendon was approximately 45 newtons (N). These results suggest that
during normal activity, a tendon in vivo is subjected
LO less than one fourth of its ultimate stress.
VISCOELASTIC BEHAVIOR (RATE
DEPENDENCY) IN TENDONS AND LIGAMENTS
Ligaments and tendons exhibit viscoelastic, or ratedependent (time-dependent), behavior under loading; their mechanical propelties change with different rates of 1041ding. v"hen ligament and tendon
specimens are subjected to increased strain rates
(loading rates), the linear portion of the stress-strain
curve becomes steeper, indicating greater stiffness
of the tissue at higher strain rates. v\lith higher
strain rates, ligaments and tendons in isolation
c',,"
store more energy, require marc force to I1.lptu
and undergo greater elongation (Kenncd.", Hawki
Willis, & Danylchuk, 1976).
Dw-ing cyclic testing of ligaments and tendo
where loads are applied and released al specific
tenmls, the stress-strain curve is displaced to
riglll along the deformation (strain) axis with ea
loading cycle, revealing the presence of a nonel
tic (plastic) cOlllponcnt; thc amount 01" perm
nent (nonrecoverable) deformation is progressiv
greater with every loading cycle. As cyclic load
progresses, the specimcn also shows an increase
clastic stiffness as a result of plastic dcfonnati
(molecular displacement). IVlicrofailure can oc
within the physiological range if frequcnt loading
imposed on an already damaged struclUrc wh
lhe stiffness has decreased.
Two standard tests that reveal the viscoelastic
or ligaments and tendons are the stress-relaxat
test and the creep test (Fig, 4~12). During a stre
relaxation test, loading is hailed safely below
linear region of the stress-strain curve and the str
is kept constant over an extended period. The str
decreases rapidly at first and then more sl()\v
\Vhen the stress-relaxation test is repeated cy
cally, the decrease in stress gradually becomes l
pronou need.
During a creep lcst,loading is halted saFely bel
the linear region of the stress-strain curve and
stress is kept constanl over an extended period. T
strain increases relatively quickly at first and th
more and more slowl.':. \Vhen this tcst is perform
cyclically. the increase in strain gradually becom
less pronounced,
The clinical application or a constant low load
the soft tissues over a prolonged period, which ta
advantage of the creep response. is a llseful tre
ment for several types of deformities. One exam
is the manipulation of a child's clubfoot by subje
ing it to constant loads by means of a plaster ca
Another example is the treatment of idiopathic s
liosis with a brace, whereby constant loads are
plied to the spinal area to elongate the soft tiss
surrounding the abnormally curved spine,
More complex viscoelastic beh41vior is observed
the entire bone-ligament-bone complex_ Antel
cnlciate ligaments (ACLs) in knee specimens tak
from 30 primates were tested in tension to failure
a slow and a fast loading rale (Noyes el aI., 1976).
, the slow loading rate (60 seconds), much slower th
that of an injury mechanism in vivo, the bony ins
tion of the ligament was 'the weakest component
the bone-ligament-bone complex.. and a tibial sp
bone-ligament-bone complex \vas nearly' the sam
These results suggest that as the loading rate is i
creased, bone shows a greater increase in streng
than does ligament.
"'"o
--'
Load relaxation
(length held constant)
Ligal1zel1t Failure and Tendon
Injury Mechanisms
Injury mechanisms arc similar for ligaments an
Time
A
tendons, therefore the following description of lig
ment injury and failure is gencrall~1 applicable
tendons. \Vhen a ligament in vivo is subjected
Creep phenomenon
(load held constant)
c
o
.~
E
loading that exceeds the physiological range. micr
failure takes place even before the yield point (P"
is reached. \"'hen Ph" is exceeded, the ligament b
gins to undergo gross failure and simultaneolls
the joint begins to displace abnormally. This d
placement can also result in damage to the su
"
<ii
o
Time
B
The viscoelasticity (rate dependency, or time dependency)
of ligaments and tendons can be demonstrated by two
Microfaiture
standard tests: the load-relaxation test and the creep test.
A, Load relaxation is demonstrated when the loading of a
specimen is halted safely below the linear region of the
load-deformation curve and the specimen is maintained at
a constant length over an extended period (i.e., the
amount of elongation is constant). The load decreases
rapidly at first (Le., during the first 6 to 8 hours of loading)
and then gradually more slOWly, but the phenomenon may
continue at a low rate for months. B, The creep response
takes place when loading of a specimen is halted safely be·
low the linear region of the load· deformation curve and
the amount of load remains constant over an extended period. The deformation increases relatively quickly at first
(within the first 6 to 8 hours of loading) but then progres~
sively more slowly. continuing at a low rate for months.
avulsion was produced. At the fast loading rate (0.6
seconds), which simulated an injury mechanism in
vivo, the ligament was the weakest component in t\Vo
thirds of the specimens tested. At the slower rate, the
load to failure decreased by 20%, and 30% less energy was stored to failure, but the stillness of the
o
2
3
4
5
6
Elongation (mm)
Progressive failure of the anterior cruciate ligament from
cadaver knee tested in tension to failure at a physiologic
strain rate. The joint was displaced 7 mm before the liga
ment failed completely. The force-elongation curve gene
ated during this experiment is correlated with various de
grees of joint displacement recorded photographically;
photos correspond to similarly numbered points on the
curve. Reprinted wirh permission from Noyes, F.R., and Graae
E.S. (976). The strengrh of the anterior cruciare ligament in h
mans and Rhesus monkeys. Age-relared and species~relared
changes. J Bone Joint Surg, 58A. 1074-1082.
350
spectivcly to (I) the load placed on the ACL duri
tests of knee joint stability performed clinical
(2) the load placed on this ligament during phys
logical activity, and (3) that imposed on the lig
ment during injury from the beginning of micr
failure to complete rupture. Microfailurc begi
even before the physiological loading range is e
ceeded and can occur throughout the physiologic
range in any given ligament. In fact, under expe
mental testing, the ultimate tensile loaels-or t
load at failure for human ACL-is between 340 a
390 N (Case Study 4-1),
Complete
failure
(3)
~
-0
"0
200
-'
o
o
2
4
3
5
6
7
8
---
Joint Displacement (mm)
ACL Failure: Failure of the ACL
Associated With High Strain and Stress
The curve produced during tensile testing of a human an-
~,.
terior cruciate ligament in vitro (20) (see Fig. 4-13) has
been converted to a load-displacement curve and divided
into three regions correlating with clinical findings: (1) the
load imposed on the anterior crudale ligament during the
anterior drawer test; (2) that placed on the ligament during physiological activity; and (3) that imposed on the liga
w
25-year·old male occasional soccer player injured
. his ACl as a result of an abnormal torque in rotation
A
of the knee. The player locked his foot on the ground
and pivoted on his lower limb to produce a high rota·
, tional torque on the knee, which increased tensile loads
the ACL,
men! from partial injury to complete rupture. It should be
on
noted that the divisions shown here represent a general-
The first region of the load-displacement curve shows
a normal physiological loading response. In the microfailure region, the increase in strain deformation leads to
high internal stress and finally a complete rupture. Experimental testing in vitro in human ACl yielded a point of
failure between 340 and 390 N.
The knee with the ACl injury will increase intra-articu-
ization. Microfailure is shown to begin toward the end of
the physiological loading region, but it may take place
well before this point in any given ligament.
rounding structures, such as the joint capsule, the
adjacent ligaments, and the blood vessels that supply [h~sc structures.
Noyes (1977) demonstrated the progressive failure of the ACL and displacement of lhe tibiofemoral
joint by applying a clinical test, the antel-iOl- drawer
'." tc;-st, to a cadaver knee up to the point of ACL failure
'AFig, 4-13), At maximum load, the joint had displaced several millimeters. The ligament \vas still in
'~o,ntinuit.y even though it had undergone extensive
~jacro- and microfailure and extensive elongation,
'In Figure 4-13, the force-elongation curve generated
during the experiment, indicaling where Illicrofail.,:. ure of the ligament began, is compared with val"ious
stages of joint displacement t-ecordcd photographically,
Correlation of the results of this test in vitro with
clinical findincrs sheds Jjcrht on the microevents that
take place in the ACL during normal daily activitv
and during injuries of variou~ degrees or s~verity. I~
Figure 4-14, the curve for the experimental study
On cadaver knees that was presented in Figure 4-J3
been converted into a load-displacement curve
divided into three regions, corresponding re~
lar joint motion, producing abnormally high stresses on
other joint structures such as cartilage, which can lead to
osteoarthritis. A deficiency in joint stability that results
from ACl impairments will increase the likelihood of ex·
periencing the "giving way" sensation or functional instability, thus affecting activities of daily living such as gait,
jogging, and squatting (Case Study fig 4-1- 1).
~
o
2
3
4
5
6
Joint Displacement (mm)
Case Study Figure 4-1-1.
7
8
Ligament injuries arc categorized clinically in
three ways according to degree of severity. Injuries
in the first category produce negligible clinical
symptoms. Some pain is felt, but no joint instability
can be detected clinically, CYCI1 though microfaillire
of the collagen fibers may have occurred.
Injuries in the second category produce severe
pain and some joint instability can be detected clinically. Progressive failure of the collagen fibers has
taken place. resulting in parlial ligament nlplure.
The strength and stiffness of the ligament may have
decreased by 50% or more, mainly because the
amount of undamaged tissue has been reduced. The
joint instability produced by the partial rupture or a
ligament is oflen masked by muscle activity, and
thus the clinical test for joint stability is usually
performed with the patient under anesthesia.
Injuries in the third categOlY produce severe pain
during the course of trauma with less pain after
injury Clinically, the joint is founelto be completely
unstable. Most collagen fibers have ruptured, but a
few Illay still be intact, giving the ligament the appearance of continuity even though it is unable to
SUppOrl any loads.
Loading of a joint that is unstable as a result of
ligament or joint capsule rupture produces abnormally high stresses on the articular cartilage. This
abnormal loading of the articular cartilage in the
knee has been correlated with carl~,' osteoarthritis in
humans and in animals.
Although injury mechanisms are generally comparable in ligaments and tendons, two additional
factors become important in tendons because of
their attachment to muscles: the amount of force
produced by contraction of the muscle to which the
tendon is attached and the cross-sectional area of
the tendon in relation to lhat of its muscle. A tendon
is subjected to increasing stress as its muscle contracls (see Fig, 6-10), When the Illuscle is maximally
contracted, the tensile stress on the tendon reaches
high levels. This slress can be increased fl.lrlher if
rapid eccentric contraction of the muscle takes
place; for example, rapid dorsiflexion of the ankle,
which does not allow for reflex relaxation of the gastrocnemius and soleus muscles, increases the tension on lhe Achilles tendon, The load imposed on
the tendon under these circumstances may exceed
the yield point. causing Achilles tendon rupture
(Case Study 4-21The strength of a muscle depends on ilS physiological cross-sectional area. The larger the crosssectional area of the muscle, the higher the magni;:'
lude of the foree produced by the contraction and
thus the greater the tensile loads transmitlcc!
through the tendon. Similarly, rhe larger the crosssectional area of the tendon, the greater the loads it
can bear. Although the maximal stress to failure for
a muscle has been difficult LO compute accurately,
such measurements ha\'c shown that the tensile
strength or a healthy tendon 111<1)-.' be mon: than
t\Vice that or its Illuscle (Elliot, 1967), This finding
is supponed clinically by the fact that muscle fUplures are more common than are ruptures through
a tendon.
Large muscles usually' have tendons with larg~
cross-sectional areas. Examples arc the quadriceps
- - - --------<0
-rend on Injuries: Achilles Tendon Injuries in
Runners. Which Result From a High Strain
Rate
.·..
A
middle-aged male marathoner engaged in a strenuous
running activity expenenced pain and a popping sensa~
tion in his posterior calf. An overuse injury is diagnosed.
The first region of the load-deformation curve ShQ\.V5 a
normal physiological toe-loading response. In the linear
region, high load is producing a higher deformation
within the tendon structure. When the Achilles tendon is
subjected to higher strain rates during frequent loading
;_ ~ycles and insufficient time is allowed for the healing
process, the result is an overuse injury. Histological studies
ot',these injuries reveal a pathological pattern described as
~'angiofibrotjc hyperplasia," which suggests a degenera" tive process. This failure in tendon remodeling frequently
,. occurs before the abrupt ruplure of the tendon. Relative
aydscu!arity, inflammatory disease, and other local factors
also contribute to midsubstance ruptures (Case Study 4·2-1),
250,00
OVERUSE
200,00
~
Toe
region
Linear
region
~
RUPTURE
yield and
failure region
150,00
u
.3 100.00
50,00
1,00 2,00 3,00 4,00 5,00 6,00 7,00
Def9rmation (mm)
Case Study Figure 4·2·1.
muscle with its patellar lendon and the triceps Slifac
muscle with its Achilles tendon. Some small muscles have tendons with large cross-sectional areas;
such as the plantaris. which is a liny muscle with a
large tendon.
did nOl differ signillcanLi~1 (Kaspcrczyk et aI., 1991).
This may be due 10 the fact lhal only lhe ACL was
taken from donors in \I,,1hom no vascular or cardiopulmonary disease and no ostcoanhritis of the
knee was found on autopsy.
Factors That Affect the
Biomechal1ical Properties of
Tendons and Ligaments
PREGNANCY AND THE POSTPARTUM PERIOD
Numerous factors alTecl I he biomcchanical properties of tendons and ligaments. The most common
are aging. pregnancy, mobilization and immobilizalion. diabeles. steroids. NSAID use. and hemodialysis. The biomcchanical properties of grafts are also
d,iscllssed because reconstruction, particularly of the
anterior and posterior knee ligaments, is common.
A common clinical observation is thc incrcasecllaxity of the tendons and ligaments in the pubic arca
during later stages of pregnancy and the postpartum period. This obsen"~llion has been confirmed in
animal studies, Rundgren (1974) found lhmlhe tensile strength of the lendons and the pubic symphysis
in rats decreased at the end of pregnancy and during the postpartum period. Stiffness of these stnlctures decreased in the early postpartum pCI'iod but
was later restored.
MOBILIZATION AND IMMOBILIZATION
MATURATION AND AGING
,
/
i':
The physical properlies of collagen and of lhe lissues it composes are closely associaled with the
number and quality of the crossMlinks within and between the collagen molecules. During maturation
(up to 20 years of age), the number and quality of
cross-links increases. resulting in increased tensile
strength of the tendon and ligament (Viidik,
Danielsen, & Oxluncl, 1982). An increase in collagen
fibril diameter is also observed (Parry et aL. 1978)
wilh high variabilily in size (range 20-180 nm)
(Strocchi. el aL. 1996) noted in lhe ~'oung «20
years), The diameler in adults (20-60 veal's) and in
lhe elderly (>60 years) decreases remarkably (120
and 110 nm) but \vith a marc even distribution.
Slrocchi et al. (1996) investigated age-related
changes in human ACL collagen fibrils and reports
an increase of libril concenlration from 68librils/mu'
in the young to 140 fibrils/mu' in the elderly. Howevel; Amiel (1991) reports Ihal lhe waleI' content
and the collagen concentration decreases significantly in the medial cruciate ligament of 2 M
, 12-, and
39-month-old rabbits,
After maturation, as aging progresses, collagen
reaches a plateau with respect to its mechanical
properties, afler which the tensile strength and stiffness of the tissue begin to decrease. This may be the
result of an increase in small collagen fibrils. Conversely, when the ACL of younger donors (average
age 30 years) was compared wilh the ACL of older
donors (average age 64.7 years), the material properties (strain. elastic modules, and maximum stress)
oF"'-';-
,
Living tissues are dynamic and change their mechanical properties in response to stress, which
leads to functional adaptation and optimal operalion of the tissue.
Like bonc, ligament and tcndon appear to remodel in response to the mechanical demands
placed on them; they become stronger and stiffer
when subjected to increased stress and weaker and
less stiff when the stress is reduced (Noyes el aL.
I 977a),
Physical training has been found to increase the
tensile strength of tendons and of the Iigament M
bone
interface (Woo et aL. 1981), Tipton and coworkers
(1970) compared the slrenglh and sliffness of medial collateral ligaments from dogs that were e~cr­
ciscd strenuously for 6 weeks with the values for ligaments from a control group of animals. The
ligamcnts of the exercised dogs were stronger and
stiffer than lhose of lhe control dogs. and the collagen fiber bundles had largel- dianlelers.
(mmobilization has been found to decrease the
tensile strength of ligaments (Newton et al., 1995;
Walsh et aL. 1993), Noyes (1977a) demonstrated a
reduction in the mechanical properties of the
bone-ligament-bone complex in knees of primates
immobilized in body casts for 8 weeks. When
lested in tension to failure, the ACLs from these
animals showed a 391'10 decrease in maximum .load
to failure and a 32% decrease in energy stored to
failul-e compared with ligaments fl'om a control
group of animals (Fig. 4-15A), The immobilized
ligaments also displayed more elongation and were
significanlly less stitT than the control spcc..:imens
(Fig, 4-158),
Amid and coworkers (1982) showed a similar de-
crease in the strength ancl stiffness or lateral collateral ligaments in rabbits immobilized for 9 weeks.
As the cross-sectional area of the specimens did 110t
change Significantly, the degeneration of mechanical properties was altributed to changes in the ligament substance itself. The tissue metabolism was
noted to increase, leading to proportionally more
immature collagen with a decrease in the amount
and quality of the cross-links between collagen molecules. Newton et al. (1995) reported that the crosssectional area of ligaments in immobilized n1bbit
knees waS 740/0 of the control valuc.
In Noves' (I977a) experiment, assessment of the
effects of a reconditioning program initiated directly aftcr the 8-\\'cek immobilization pel"iod
demonstraLcd that considerable time was needed
1'01' the immobilized ligaments to regain their former strength and stiffncss. After 5 months, the reconditioned ligaments still showed considerably
less stiffness and 20% less strength than did ligaments from control animals. At 12 months, the reconditioned ligaments had strength and stiffness
values comparable to those of control group liga-
ments (Fig. 4-1 SA). Woo et "I. (1987) found that the
stress-strain characteristics after rcmobilization return to normal but that the energy-absorbing capabilities or the bone-ligament complex improved but
did not return to normal.
DIABETES MELLITUS
The term diabetes rcfcrs to disorders characterized
by excessive urine excretion. Diabetes mellitus is a
metabolic disorder in which the ability to oxidize
carbohydrates is more or less completely lost. This
is usually caused by pancreas insufficiency and a
disturbance of the normal insulin mechanism, rcsulting in hyperglycemia, glycosuria, and polyuria.
Diabetes mellitus is known to cause musculoskeletal disorders. Diabetics compared with nondiabetics
show higher rates of tendon contracture (29 \IS. 9%),
tenosynovitis (59 vs. 7%), joint stillness (40 vs. 9%),
and capsulitis (16 \'s. 1%). Diabetes also causes osteoporosis (Calvallo et aI., 1991; Lancaster et aI.,
1994).
Duquette (1996) examined the effects of diabetes
on the properties of the collateral knee ligament in
rats. The tissue elastic properties did not differ between the diabetic and the control group. The vis-
100
60
-"
0,2
~~
l'l E
c ~
"~·x
E
"'"
60
40
Q..E
20
o
Maximum load
10 failure
Energy slored
10 failure
--
,.- "
"
""
"
A'
.-------
Immobilized
--~8 weeks)
I
\
\
\
Elongation (mm)
B
A
A, Maximal load to failure and energy stored to failure for
primate anterior cruciate ligaments tested in tension to faHure. Values are shown as a percentage of control values for
three groups of experimental animals: (1) those immobilized
in body casts for 8 weeks; (2) those immobilized for 8 weeks
and given a reconditioning program for 5 months; and
(3) those immobilized for 8 weeks and given a reconditioning
I
program for 12 months. 8. Compared with controls. ligaments immobilized for 8 weeks were significantly less stiff (as
indicated by the slope of the curve) and underwent greater
elongation. Adapred from Noyes. FR. 0977(1). Functional proper~
ties of knee ligaments and alterations induced by immobilization.
(lin Onhop, 123,2/0-242.
component of the tissue response, however,
\vas increased in the hy'perglycemic group. Insulin
therapy seems to lessen such alterations. Lancaster,
ct al. (1994) examined the changes in the mechani~
cal properties of the patellar tendon in diabetic dogs.
The results showed the stiffness of the canine patellar tendon-tibia complex in a phy'sioiogical range of
loading was 13% greater than in the control group.
There was no difference in the strength of the tendon
between the groups, but the mode of failure was dif~
ferent. In the control group, failure was caused by
substance and avulsion failure, whereas failure of
the tendon in the diabetic group was caused by tensile fractures of the patella (Lancaster et al" 1994).
COllS
STEROIDS
Corticosteroids, when applied immediately after injury, may' cause significant impairment of the bio~
mechanical and histological properties in liga~
ments. Corticosteroids also are known to inhibit
collagen synthesis in vitro (\'Valsh et aL, 1995). \Viggins et a!. (1994) described these results in rabbits
and implied that an acute injured ligament treated
with corticosteroid injections may not \vithstand
the mechanical loads of an earl)!, vigorous rehabilitation. Noyes et a!. (1977b) reported decreased ligament stiffness, failure load, and energy absorption
in monkey ligaments after injection of long-acting
corticosteroids. These findings were tin1e- and
dosage-dependent. After application of a dosage
that was approximately 10 times an equivalent human dose. only! minimal changes were found after 6
\veeks, but after IS weeks the maximum failure load
(20%), energy absorption prior to failure (11%), and
stiffness (11 %) decreased significantly. After
application of a dosage equivalent to the human
dose, the maximum failure load (9%) and the enabsorption (8%) decreased significantly.
Campbell ct al. (1996), however, showed that a
single injection of long-acting corticosteroid does
not cause histological differences in rats with acute
injured ligaments as compared \vith rats with acute
ligament injury and no injection of corticosteroids.
lv1cchanical testing showed no significant difference
in ultimate load or ultimate stress in the two
groups. Oxlund et a!. (1980) reports that local injections of corticosteroids every 3 days for 24 days increase the tensile strength and maxin1um load stiff~
ness of muscle tendons but decrease the strength of
the bone attachments of ligaments.
Laborator).' investigations established the presence of estrogen receptors in hun1an ACLs. Liu et al.
(1997) reports that physiological levels of estrogen
reduce the collagen production by 40(;1(; and at phar
macological levels of estrogen, collagen production
is decreased by' more than 50%, Estrogen fluctua
tions may alter ligament metabolism and ma
change the composition of ligament, rendering i
more susceptible to injury.
NONSTEROIDAL ANTI-INFLAMMATORY DRUGS
NSAIDs (which include aspirin, acetaminophen, and
indomethacin) are frequently used in the treatmen
of various painful conditions of the musculoskeleta
system. NSAIDs are also widely used in the treat
ment of soft tissue injuries such as inflammatOl)' dis
orders and partial ruptures of tendons and liga
ments. Vogel (1977) found that treatment wit
indomethacin resulted in increased tensile strengt
in rat tail tendons. An increase in the proportion o
insoluble collagen and in the total collagen conten
also was observed. Ohkawa (1982) found increase
tensile strength in the periodontium of rats after in
domethacin treatment. Carlstedt and associate
(1986a, 1986b) found that indomethacin treatmen
increased the tensile strength in developing an
healing plantaris longus tendons in the rabbit an
noted that the mechanism for this increase wa
probably an increased cross-linkage of collagen mol
ecules. These animal studies suggest that short~tern
administration of NSAlDs would not be deleteriou
for tendon healing but instead \vould increase th
rate of biomechanical restoration of the tissue.
HEMODIALYSIS
Tendinous failure resulting from chronic renal failur
does occur, \vith tendon rupture reaching 36% amon
individuals receiving hemodialysis. Hyperlaxity o
tendons and ligaments \vas found in 74%, patella
tendon elongation in 49%, and articular hypermobil
ity in 51 % of individuals receiving long-term he
modialysis (Rillo et a!., 1991). Dialysis-related amy
loidosis may cause the deposition of amyloid in th
synovium of tendons. The major constituent of th
amyloid fibrils is the beta 2-microglobulin (Morita e
a!., 1995; Honda et a!., 1990; Bardin et a!., 1985).
GRAFTS
Reconstruction of torn ligaments, especially or th
anterior and posterior cruciate ligament, is now
frequent procedure. The need for reconstruction
related to age, activity level, and associated injurie
Grafts derived from different individuals of the
same species ar~ called allografts; grafts derived
from the same individual are called aUlograhs. Allografr tissue preservation is done through freezcdrying and low·dose irrndiation to reduce rales of
rejection ancl infection and to limit effects on the
structural properties. Bone-patellar. tendon-bone,
and Achilles tendon are llsually L1sed as allograft tis-
sue, whereas the central tissue of the patellar tendon is commonly llsed as autograft tisslie.
Shino ct al. (1995) L1sed fresh-frozen allogenic
Achilles, tibialis anterior or posterior. and peroneus
longus or brevis tendons for ACL reconstruction in
humans. Specimens were procured during secondlook anhroscop~'. Several years after reconstruction.
the allografts had collagen fibril profiles that did not
resemble normal tendon grafts or normal ACL.
Strocchi et al. (1992) L1sed patellar tendons that had
been autograftcd to reconstn1C[ [onl ACLs. Follow-up
biopsies were performed 6, 12. and 24 months after
surgel}'. During this time, the aUlOgraft underwent
considerable changes. and after 24 months the autograft had the appearance of normal ligament tissue.
Strocchi suggested that the patellar tendon autograft
is a valid functional ACL substitution for patients who
desire to perform normal mechanical activity.
Corselli el al. (1996) reports thal replacement tis-
sue undergoes extensive biological remodeling and
incorporation. However. even a fully incorporated
graft will never duplicate the native ACL but works
instead as a check reign that increases the knee
function_ Tohayama et al. (1996) stated that the
graft elongation at the time of implantation influences the long-term outcome of ACL rcconstruc·
tions. at least in the canine model. They compared
those cases where the graft elongation behavior was
within the 95% confidence limit of normal ACL
(group 1) with those cases where the graft elongation behavior was more than lhe 95(1'0 confidence
limit of the normal ACL (group 2). Group 2 had sig-
nificantly less inner stiffness of the graft than did
group 1. Group 2 showed a significantly increased
anteroposterior laxity. but there was no difference
in ultimate failure load and absorbed energy.
substantial proportion of elastin. which lends these
structures their great elasticity.
2 The arrangement of the collagen fibers is
nearly parallel in tendons. equipping them to with-
stand high unidirectional loads_ The less parallel
arrangement of the collagen fibers in ligaments al~
lows these structures to sustain predominant tensile
stresses in one direction and smaller stresses in
other directions.
,;:~~)': At the insertion of ligament and tendon into
stiffer bone. the gradual change from a morc fibrous
[0 a morc bony rnaterial results in a dccreased stress
concentration effect.
~ ~ Tendons and ligaments undergo deformation
before failure. ,,""hen the ultimate tensile strength of
these structures is surpassed. complete failure DC·
curs rapidl~!. and their load~bearing ability is sub~
stantially decreased.
5. Studies suggest that during nonllal activity. a
tendon in vivo is subjected [0 less than one fourth of
its ultimate stress.
'/'~:(: Injul"\! mechanisms in a tendon are influenced
by"{he ~lm~unt of force produced by the contraction
the muscle to which the tendon is attached and
the cross-sectional area or the tendon in relation to
that of ilS muscle.
or
7 i The biomechanical behavior of ligaments and
tendons is viscoelastic. or rate·dependent. so thal
these structures display an increase in strength and
stiffness with an increased loading rate.
S An additional effect of rate dependency is the
slow deformation, or creep. that occurs when tendons and ligamcnts are subjected to a constant low
load over an extended period; stress relaxation takes
place when these structures sustain a constant elongation over time.
/{gi,::':
Anina
results in a decline in the mechanical
,/),,,
:;;.
b
properties of tendons and ligaments, th~1t is, their
strength, stiffness, and abilily to withsland defor-
mation.
10
Pregnancy, immobilization. diabetes. steroids.
NSAlD Lise, and hemodialysis affect the biomechan-
Summary
f~1-~Tendons and extremity ligaments are composed
largely of collagen, whose mechanical stability gives
lhese structures their characteristic strength and
flexibility. The ligaments flava of the spine have a
ical properties of ligaments and tendons.
11
Allografts and autografts are useful in ligament
reconstruction but material propenies do not re'turn completel~" to normal levels.
12) Ligaments and tendons remodel in response to
the mechanical demands placed on them_
REFERENCES
:\micl. D.. Kuipl'l'. S.D.. \\'all"C('. C.D .. Harwood. E V'lndl'berg. J.S. (1991). :\gt:.rdall'd properties of llledi.d (;olbt·
l.:ral Iig~lIncnt and anterior cruciall' ligament: ..\ morpho·
logic :lnd coll'lgen maturation study in thl' r'lbbit. J
Gerem/ol. 46(.J). BI56-BI65.
Amid. D•• Fr,lIlk, c.. I-Iar\\'ood. E. l.'1 OIL (198-1). Tendou~ and
ligaments: A morphological and biochl'lllical comp~l'ison.
lOt/hop nt's. 1.157.
:\mit'1. D., Woo. S.L.Y.. Harwood. F.L.. l't a!. (1982). Thl' dfl.'ct
of immobilization on colbgen lUl'IlO\"l'J" in t:onlll,,'cli\'c tis·
suc. i\ biochl..'rnical-biomechanical C01Tl'btion. Ac/o Or.
!hop Scalld, 53, 323.
Bardin. T.. Kuntz, S., Zingr.d"f, J.. Voisin. ~'I'I.C., ZelmaI', A.
(1985). Syno\'ial amyloidosis in patients undergoing long:.
terlll hl..'lllodialysis . .·\J'/l1l"it;., Rheum, 28(9}. 1052-1058.
Campbell, R.B .. Wiggins. \-1. E.. Cannislra. L.;\L F"dalc. P.P..
Akclm.m, E. (1996). Influencc of stcroid injc(:tion in li~:lment hc~t1ing in thl..' r.lI. Cli" OnllOp. 332. i.J1-253.
Carlstcdl, C.:\. (1987). Mt:chanical and cht.'f1li(~11 f:u;tors in
[('ndon healing: Effects of indolTlcthacin and sun.!cn· in
th..: 1'00hbil. Aew On/WI' Scawl SIIPI,I, IN.
~ .
, Carlstcdt, C.:\ .. Madson. K.. &. Wn:dmark. T. (l9Soa). The illflllencc of indolTlethacin on collagen synthi.'sis durin!! lell'
([cm healing in lhi.' rnbhit. Pros/ag!(/I/{!i,ls. 32. 353. C~Il'lstcdl, C.A .. ~bdsOTl. K" & W!'cdrnark. 1". (1986b). Till' influent:l.: of indtlllll.'thacin on lendon hl'aling. A biolJlcchall'
itnl nnd biochcmical sllJd~'. ..lFell OrtllOl~1 7i'allll/{/ So!'!.;.
105. 332.
'
(:11'\,(\110. A.. Ordonez. M.A .. Garcia. H.. T;tpia. J.C .. Bcbv:\c,
1.. V:tlC:.'llCi;l. J., r..lorl'ira. ~','l. (1991). Adult llon·insll1in-dt:pcndl'lll diahetic: Limitation of <lJ'licular Illobi1itv ~\lld soft
tissue involvcment. Rev lIed Chil,119(9). 10Io-1()2l.
Cooper. R.R .. ~ Misol. S. (1970). Tendon .Ind lig:lIllCIH inserlion. ;\ light and electron microscopic sllLdv. J B{J!/~' Joiul
SU,,!;. 52;\, I.
.
Corsl.'tti. J.R .. J'lckson. D.W. (1996). Fnilul'c of :lIl1Crior nud·
~IlC Iig<lll1l'llt rl'conslruction: The biologic basis. Clill Or.
fhop. 325. .J2--t9.
Diamant. J .. Kdlcr. r\ .• Bael', ..\., Lill, M.. :\lTid~l'. R.G. (1972).
Colhlgen: Ultraslructure and its relations ~ 10 mechanic,,1
propl..'ni<.'s as a function of aging. Prot· R Soc Loud
iBio/}, /80, 293.
DuquclIc. J.J .. Grigg, P.. &. Hoffman. A.H. (1996). The effect
(If diabetes on the viscoelastic propcrtil's of r~\[ kncl' li!'!H·
ments. J Biolllt.!ch Ellg. 118(·0. 557-56.J.
Elliot!. O.H. (1967). Tht: biomcchanical propl'rties of ll~lldoll
in relation 10 muscular strcn"lh . .-\111/ PIn's ,\led 9 1
Fillon·Jne.:kson, S. (1965), Antc~cdcnt pha~cs'in ;lla'tJ"i~ formalion. In 5tl'/lt"lll1"1..' alld PI/llelioll OJ"COIIIIt:ctiw {jllll Skelewi Tis.~lIes (p. 277). London: ButtCl'worth.
Fun~, .Y.C.~. (1981). Biomeclumics: ;\kcllflllical Propr:f,i.:s o{
Llvl1Ig T,ssue.'s (p. 222). Nc\\' York: Springer· Veri:.!!.
Fung..r.C.B. (967). Elasticity of soft lissucs in sim-pk don·
g'lllon. :Im J Phvsiol, 213. 1532.
Fung. Y.C.B. (1971). Strl.'ss-str.lin-hislo/,v rdations of sof! lis·
sues in simplc elongation. In Y.C_ Fu·ng. N. PelTOnl', & t\'1.
l~nlikcr (Eds.). Biolllechallics: Its FOlll/dations awl Ob;ec. IlvcS (pp. 181-20$). Englewood CliiTs. NJ: Prentice-Hall.
HIrsch, C. (1974). Tensile properties during tendon healinl!.
..lCla Or/hop Scellld. SIIPl'l 153.
-
Honda. K.. I-lara. M., Ogura, Y.. Nilll.:i. 1-1 •• i\lirnura, N. (I990
~~t<l-~-microgl()hulin ~llnyloidosis in hemodialysis pa
t~l·ntS: An autopsy study of intcrH·rtt:hr.t1 disks and POSlC
nor hg<lOll·lltS . ..\eUl Patlwl lpII . .JO( II). S20-826.
Kaspl'I"czyk. W. J.. Rosoch:l. S .. BO$t:h. U .. Ocsh~rn. J.1.. Tsch
crnl·. I-I. (1991). Ag(' .\CtiviIY .Hld stl'<.'ngth of knee.: li!:.!
ments. U,,{alkhimrr,. 94(7).372-375.
Kassel". J. (Ed.) (19961. Onlwp/li.'dic I\lI00dccly,i.' Update
f/?IIll! S{lfdy Syllal}f/,~, P'.ll'k Ridgl'. IL: American Acadl'Ill
ot Onhopatdic Slll"~c.:ons.
Kcar, i\l.. & Smith. R.N. (1975), A flll'thnd for l'l'corJin u W
don stnlin in sheep during locomotioll, Aew On/lOp S:flll
-16. 896.
Kl'llIll'dy, J.C., l-bwkins. RJ .. Willis. R.B .. d :.1. (1976). Ten
sion Sllu.!i(·S of humall knec ligaments. Yield point, ult
matc failure, and disruption of thl.: cl'tlciatc and tibbl co
"Heral ligaments. 1 BOlle.' loillT S/f/'~. 58:\. 350.
Lancaster. R.L., Haut. R.C.. DeCamp. C.E. (199.J). Chan2es
thl' lllech:luic,,1 propenil's of p.ltdl.ll· tl'lldOll pr..:par.~tio
of spolHancously dii:lbl'lic do,gs tIIH!l'r 10nll·tC'I"1ll insuli
th~rapy. J Biomf!cil. 27(8). 110.1-1108.
Lill. S.H .. AI-Shaikh. R...\., Panossi~lll. \1.. Fil1{·rman. G.A
Lane, J.M. (1997). Estrogen ;tlTccts l!Ie cellular Illctabo
lisll.l of .lh(.' .aolt'rior t:l'l1ciall' ligaml'lll. ,\ potential expl
n:tllnn lor lemalL' ~llhlt.'lic injury..'\111 1 Sports ,I/i.'d. 25(5
71J4-709.
:\-lorita. H.. Shinzaalo. T.. Cai. Z.. Dadd. G.. i\lizulani A
Habuchi, H.. Ito, :\1.. Asai, J. 1.'1 ;d. (1995). Basic fibrol;la
growth faClor-hepatin sulphate complex in the huma
dia1:o'sis-relalcd amyloidosis, Vir"hol\'s Arch. ';27(4
395-400.
Nachclllson. A.L .. & E\'ans, J.I-I. (1968), Some mechanic
propertit:s of tht: third hUl1wn lumb;\[' intcrlaminar lig
men! (ligalll(.'ntllm lhvlllll). J BicHIIIN,:". I. 211-220.
Nl·wton. P.O., \Voo. S.L.. ~ladenrw, D.A .. Akeson. \V.H
(1995). Immohilization of tli..: knee:.: joint alters Ihl' m
dwnic~d and uilrastrucllll',ll propenil's of the r.lhbit anle
rior crucialC Iigaml'llt. J Orthop UI.:.... 13(2). 191-200.
Noyl's. F.R .. t:l al. (l977a). Funt:lional propl'nics of knee li
:1I11,;:-Uls and alterations induced b\' immobilization. Cl~
On/lOp. 123, 210-1-12.
.
Noyes. F.R .. Grood. E.S.. Nussbaum N.S .. Cooper. S.M
(1977b). Effect of intra<lrlicuhlr corticosleroids on li!:.!
mCIH properties: A biollll'chiHlic;1! and hislOlo!'!ical stl~
in Rhesus knees. Clill art/lOp. 123. 197-209.
~
No)'l's. F.R .. and Grood, E.S. (1976). Tht: sm:ngth of the antl
riot' cruciate ligament in hUIlIi.111S and Rhl'SllS monkl'\'
r\gc-rdaled and specil's.rl'lated chan!!,:;.. J BOlli: loint SI/·,
58..\, 1074-1082.
~
Ohkawa, S. (1982). £ffl'ns of orthodolltic forces and anti· i
n<lmn~a(or~' drugs on the !llct:hallical sll'l:~ngth of the per
odoll II II III 111 Ihl' ral mandibular first molar. ..\/1/ J artho
8/, 498-502.
Oxlund, H. (1980). The influence of a lot:~ll injt:t:lion of cort
solon the Illcchanic,,1 prolll'nies of tendons ;lnd lit!amelH
and thl.' indirect effeci 011 skin . ..\eta Or//top St:tllld, 51(2
231-138_
P.llTY, D.:\.D., Barnes. G.R.G ..•lIlel Cr;Ii!!, :\.5. (1978).:\ com
p"rison of the size distribution of 1.'~lI;l!.!cn fibrils in co
nective tissucs as a function of age anl possible rdatio
between fibril size ;\Ild llll'chanicO'1 propcrlil's. Proc R So
Lrmd. 203. 305-321.
Prockop. D.J .. Ii: Guz.man, N...\. (1977). Collagl.:1l dist.'asi,.'s ;\ltd
the bios,\'lllhcsis of collagen. /-Iosp Prw.:t, Dec, 61-68.
Rarnachalldr<\ll. G.N. (1963). Molecular strllC!lln: of collagen.
lilt Rev COllllcet TisslU' Uc.~·, /, 127.
Rich, A., Crick, F.H.C. (1961), The molecuhll' structure of col·
lagen. J .Hol Bioi, 3. 483.
Rillo. D.L. Babini. S.M., B<lsn;lk. A., Wrlincr. E., BaHwclwl1.
E.. Cocco. J.:\. (1991). Tendinous .md Iig:Ulll'lltoliS hyperlaxity in patients reechoing long-Icrm hemodialysis. J
Rht'lIl1ullof. /8(8), 1227-1231.
Runc.lgrcn, A. (1974). Physical properties of conncctin: Lissuc
as influenced by single and n::pc<I!L.'d pr~gn;\ncics ill thl.'
r;IL .·kltl Pilysiol Scam!, Sl/ppl.. 417.
Shino, K., Oakt-s, B.W.• I-Ioribe. S.. Nnkaw, K.. Nakamurn. N.
(1995). Collagen fibril popuhltions in human ;:Ulterior cru·
ciate lig;unent allografts. Electron micfoscopic analysis.
A/1/ } Sport.~· Mal, 23(2), 203-208.
Simon, S.R. (1994). Orthopedic Basic SciclI"'~. Rosemont, IL:
MDS,
Snell. R.S. (198-1). Clillicld al/{! FI/Ilctiomd Histology {or ,\It,d·
icaf SIUt/CIIIS. Boslon: Lillk" Brown.
Stfocchi, R., Dl' P<lsquale. v., Facchini. ,\ .. Rnsp,ulli, M ..
Zaffagnini, 5 .. ~'larcacci, ~1. t 1996). Agc:·rcbted changes in
human anterior cruciatt: ligament (ACL) collagen fibrils.
Ira! ..lWI/ E",bryui. /01(-1).213-120.
Slrocchi, R.. DePasqu:lIL', V.. Guizzardi, S., ~I:lrcacci, M..
Ruggeri. A. (1992). Ultr.\slructural modificalions of patdlar tendon fibres lIscd as 'Ulterior cruci.\te ligament (ACI..)
repbcclllcnL/uJ! } Allat EII/bryo!, 97(4). 221-228.
Tipton, C. =\<1. , James, S.L., Mcrgner, W., Cl al. (1970). Influ·
encc of e-xcn::ise on slrl.:ngth of medial (ollatcralliganH:nls
of dogs. Alii) Physio!, 2/8. 894.
Tohayama. 1-1 .• Beynnon. B. D.. Johnson, R.J., Renstrom, P.A ..
Arms S.W. (1996). The effect of <lnlerior cruciate Jig.lInent
graft c1ong'llion at the lime of implantation on the biolllechanical behavior of the graft and knee. ,.\/1/ j Sports .\lcd,
24(5),608-614,
Viidik, A. (1968). EJ<lslicil)' <IIul lensile strenglh of Ihe antl.'·
rior cruci<llc ligament in rabbils :lS influcnced by Iraining.
Acta Ph\'sio! Scallll, 7~, 372.
Viidik, A. '(1973). Functional propertics
collagenous tissues. IlIr Rei' COllllcct Tissue Res, 6. 127.
or
Viidik, A., D'lllidscn, C.e.. Oxlund, H. (1982). Fourlh tnlem,ltional Congress of Biorhl'oJogy Symposiulll on Mcc.:hHnical
Propcnics of Living Tisslles: On fundamental and plll:llOnlcnological models, SlrllC(llrc and mcch:lJlical properties of
collagen, elastic and glycosaminoglyc'1I1 complexes. Biorll/:o!o,~y.
19,437.
Vogel. 1·1.e. (I 97i). ~'kcll<:lnh.:al and chl'mic:d properties of nil'·
iOlls conneclive lissue org.tns in rats as iniluenced by nonsleroidaJ anlirheum:ltic drugs. COI11/l!ct Tisslle! Res. 5, 91.
Walsh, W.R., Wiggins. r-,·1.E .. Fadak P.O., Ehrlich, !\·I.G.
(1995). Errc:cts of dd;tyed steroid injection on ligamenl
healing using a rabbit medial colialen,1 ligament model.
BilJllwt..:,-;a!s. 16( I 1), 905-910.
Walsh,S .. Fr'lIlk. c.. Shrive. N., Han, D. (1993). Knee immobiliz;Hion inhibits biolllcchanic<llmatllnHioll of the r:\bbit
medinl collateral lig:ll11l:111. Clill Orrhop. 297, 253-261.
\Vhite, A., H;:lndler, P., & Smith, E.L. (1964). Prillciples o{Bio.
cI/!:/IIistry. Ncw York: McGraw-Hill.
\Viggir\,o;, [VI. E., Fadale, P.O., B<lITach, H., Ehrilich, ,,"I.G.,
Walsh. W.R. (1994). Healing ch<lr;tClcrislics of a type I col·
Iag('1l01lS SlnlClUre treJh:d wilh corticosteroids. .-\1/1 j
SPOI"/S ,\kd, 22(2), 279-188.
Woo, S.L.Y.....\n, K.N., Arnoczky, D.\I.\1.. Fithian, D., & ~ly('rs
B. (1994). Anatomy, biology, and biomechanics of the lendon, ligaml'nl, ;Iud m..:niscus. In S.R. Simon (Ed.). Or·
rJ/Of/at'die Basic SCit'IICt' (p. 51). Ros\,'monl. Il: .-\AOS.
Woo, S.L.Y., Gomez, \-1.,.\.. Siles, T.J., Cl :11. (1987). Thc bin·
mechanical .llld morphological clwngcs in the Illcdi:d (01latcr,tl ligamelll of the rabbit after imlT1obilization (lnd rc[T\obiliz;llion. ) !3lme )0;1// SlIrg, 69..1(8),1200-1211.
\Von, S.L.Y., Gomez, M.A., Amid, D., Ritl('r, ~·l.A., Gdbt.'rman,
R.I,!.. Akcson, W.H. (1981). The effects of e.\crcise on the
biomcchanical and biochemical propcnies of swine digiwl
flc:.\or leudons. } B;ol1l~!c.:h Ellg, 103, 51.
Woo. S.L.Y. (1988). Lig;\mellt, tendon. :HH.l join! capsule insertions tu bone. In S.L.Y. Woo, & J. Buckwalter(Eds.), Illjllly (II/(/ Repair of th~ ,HlISCU!oskduli So{t Ti...... uc:s (pp.
133-166). Park Ridge. II. :\meric.w Academy of Or·
Ihop<:ledi(..· Surgeons.
,,\',
" ST~UC,;,UR~
;;'
,~
',,'-
Nutrition. mct.'lbolism
CEL(:FIBROBJ-ST't;\, !
.~,
~'\.\,
\},o'
Synthesize
Ex(r!:lcclJuI3r Mmix
81
l
"" EXTRACEllltLAR ."
-\
STRONG AND
FlEXIBlES,
-{
The
-Collagen type J: suPPOrt tensile loads
mcc~anjcal
response
is dependent on the
rate at which thl! loads
arc applied
-Ground $ubsL.,ncc
Gel.like matcrialfSt.abilizc
collagenous skclC!ton
Resistance of
high tensile
forces with
limited elongation
'El:min
pS;.;~"NEUAAl
Nm:ORK'
M~chjlnoreccp[ors
, :.~\:' ,-
·Proprioception
·Sensorimotor Mechanism
'~\,
'-
',;:'-c:,,;,,'
FLOW
CHA~T
4-1
Common structure and mechanical properties of tendons and ligaments. *
*This flow chart is designed for classroom
Of
group di~cussion. Flow chart is not meant to be exhaustive.
..
..~.~."" ,~,."""".~~
..•""",,,,,,,,,,,,,,,,,,'",,"_~,_,,~,,,,.,,-"",",,,,,,,,,,,~,_,,,,,!.'<=J ••""......"
~.
;~
;.'
,(
ci
,~
y,
"•
x0
0
~
9
"
0
•~
.~
to
w
"-
Ci
-5
;;;
u
·c
m
.c
u
w
E
c
"0
m
~
0
10
~
~
0
"0
c
~
'\:z::
):Sf
<5c
.to
0
,
0
E
0
•
~.
C
0
.~
~
'C
c.
0
0
'<;"
E
0
~
:::
u
.2
•m
~
C
."
,,:,;iO{:';
i,;vi
0
-"~/,<
~
';:>'i',·o-
;:{;y;,
,
/
.-
to
0
-5
•
0
;0
i'o
~
s;;:e
6
,~:<t
;;i
~;;
';~
I'
'r."
l:
The Cell andfor eXlracel1ubr matrix components
,
1?
I
t'
'{
Mechanical Caused InjurJes
Injuries in association with local strain type
and/or level of stress
;,
i{
~:
:{I
'"
~I
j
t
),I!
Stre5s
+- Str:lin
J
l
Microtr.lumJS that
exceed the reparative
procc5S
...
)l
t
_._-
Stress
-Thermal injuries
Tendon
'Gun injurie5
'Inherit injuries
'Paratcnonilis
·L1cer3.tions
'Corticosleroids
·Tcndinosis
'Bone fr.lCIUre5
'Endocrinologic and
metabolic discases:
diabetes
-Tendinitis
'Colli5iol1s
Tendon's Jtuchment
-Rheumatoid arthritis
'ApophY5it is
-Congenital diseases
-EnthcsopJtics
'Inf~ctions
\)
i
RUPTURE
,
"1
:l);
TOlal/PJrti;a1
,if
FLOW CHART 4·3
Tendon injuries. Clinical examples. *
"This flow chart is designed fOf cla5sroom Of group discussion. Flow chart is not meant to be exhaustive.
'/01
. .".._.
-'--~---'-'-=--
J
Function
Structure
ViSCOElASTIC.
MECHANICAC
PROPERTY
FLOW CHART 4·4
~This
Ligaments structure and mechanical properties.* (PG, Proteoglycan)
flow chart is designed for classroom or group discussion. Flow chart is not meant to be exhaustive.
.:;
The Cell and/or cxtracellular matrix componcnts
Injurics in
a~sociation
Injuries in association with local strain type
and/or Icvel of stress
with diseases
*1
Dccre.lsc in mechanical
properties
),
~
i:
+-
~i!
;J
Stress
Stress
·Endocrinologic ;lOd
metJbolic diseases:
Diabetes Mellitus
'Collisions
-l~cerations
Tot~I/PJrtial
I
t
-Gun injuries
RUPTURE
'1
unit area
'Bone fnlCtureS
''.t.,."I
'FLOW CHART 4·5
t Amount of load per
'Iatrogenic Injuries:
Corticosteroids
I
I
-+ Strain
t Sinin r.ltC
Microtraumas that
exceed the reparative
process
Ligament failure. Clinical examples.'"
-This flow chart is designed for classroom or group discussion. Flow chart is not meant to be exhaustive.
Biomechanics of
Peripheral Nerves and
Spinal Nerve Roots
Bjorn Rydevik, Goran Lundborg, Kje/l O/marker. Robert R Myers
Introduction
Anatomy and Physiology of Peripheral Nerves
The Nerve Fibers: Structure and Function
Intraneural Connective Tissue oi Peripheral Nerves
The Microvascular System of Peripheral Nerves
Anatomy and Physiology of Spinal Nerve Roots
Microscopic Anatomy of Spinal Nerve Roots
Membranous Coverings of Spinal Nerve ROOlS
The Microvascular System of Spinal Nerve Roots
Biomechanical Behavior of Peripheral Nerves
Stretching (Tensile) Injuries of Peripheral Nerves
Compression Injuries of Peripheral Nerves
Critical Pressure levels
Mode of Pressure Application
Mechanical Aspects of Nerve Compression
Duration of Pressure Versus Pressure level
Biomechanical Behavior of Spinal Nerve Roots
Experimental Compression of Spinal Nerve Roots
Onset Rate of Compression
Multiple levels 01 Spinal Nerve Root Compression
Chronic Nerve Root Compression in Experimental Models
Summary
References
Flow Charts
•
._~ ..
_-...
_-~
Introduction
The nervouS s)'stcm scn'cs as the bod.:,/s control center and communications nct\vork. As slIch, it has
three broad roles: it senses changes in the bod)! and
in the external environment, it interprets these
changes, and it responds to this interpretation b)·
initiating action in the form of muscle contraction
or gland secretion.
For descriptive purposes, the nen'OllS s)'stem can
be divided into two parts: the central nervous system, consisting of the brain and spinal cord, and the
peripheral nervous system, composed of the various
processes that extend from the brain and
spinal corel. These peripheral nerve processes proinput to the central nervous system from senreceptors in skin, joints, muscles, tendons, viscera, and sense organs and provide output from it to
effectors (muscles and glands). The peripheral nervous system includes 12 pairs of cranial nen'es and
branches and 31 pairs of spinal nerves and
their branches (Fig. 5- lil). These branches arc
called peripheral nerveS.
Each spinal nerve is connected to the spinal cord
a posterior (dorsal) root and an anterior
root, which unite to form the spinal nerve
intervertebral foramen (Fig. 5-1, B-D). The
noste,ric)r roots contain fibers of sensory' neurons
conducting sensory information from recepin the skin, muscles, tendons, and joints to the
celHr.al nervous system) and the anterior roots con~
mainly fibers of motor neurons (those that conimpulses from the central nen'ous system to
targets such as muscle fibers).
Sh,ortlv after the spinal nerves leave their interforamina, they divide into two main
the dorsal rami, which innervate the musand skin of the head, neck, and back. and the
generally.' larger and more important ventral rami,
innervate the ventral and lateral parts of
structures as well as the upper and lower ex~
,,.,'m;t;,>< Except in the thoracic region, the ventral
do not run directly to the structures that they'
innervate but first form interlacing networks, or
with adjacent nen·'es (Fig. 5-1A).
This chapter focuses on both the peripheral nerves
spinal nerve roots, which cOlltain not only nerve
fibers but also connective tissue elements and vascular structures that encompass the nerve fibers. The
possess some special anatomical properties
that may sen/e to protect the nente from mechanical
damage, for instance, stretching (tension) and com-
pression. In this chapter, the basic microanatomy
the peripheral nentes and the spinal nCIYC roots
revicwed \vith special reference to thesc buil
mechanisms of protection. The mechanical behav
of peripheral nen'es that are subjected to tension
compression is also described in some detail.
Anatomy and Physiology
Peripheral Nerves
of
The peripheral nervcs are complcx composite str
tures consisting of nen'e fibers, connective tiss
and blood vessels. Because the three tissue cleme
that make up these nen'es react to trauma in dif
ent ways and may each pia)' distinct roles in
functional deterioration of the nerve after inju
each element is described separatel)'.
THE NERVE FIBERS: STRUCTURE AND
FUNCTION
The term nerve fiber refers to the elongated pn)c
(axon) extending from the nerve cell bod)' along w
its myelin sheath and Schwann cells (Figs. 5~2
5-3). The nerve fibers of sensory neurons cond
impulses frolll the skin, skeletal I11uscles, and jo
to the central nervous svstem. The nerve fibers of
motor neurons conve)' impulses from the cen
nen'ous system to the skeletal muscles, causing m
cle contraction. (A detailed description of the
chanics of muscle contraction is given in Chapter
The nen'e fibers not only transmit impulses
also sen'e as an anatomical connection between
nerve cell body' and its end organs. This connect
is maintained by axonal transport systems, thro
which various substances synthesized within
cell body (e.g., proteins) are transported from
cell body to the periphel)' and in the opposite di
tion. The axonal transport takes place at speeds t
Vat)' from approxilnatcly 1 to approximatel)"
mm per day.
Most aXO!lS of the peripheral nervous system
surrounded by multilayered, segmented coveri
known as myelin sheaths (Fig. 5-3). Fibers with
covering are said to be myelinated, whereas th
without it (mainly small sensol)' fibers conduct
impulses for pain from the skin) are unmyelina
The myelin sheath of the axons of the periph
nerves is produced b).' flattened cells called Schw
cells arranged along the axon (Fig. 5-3). A sheat
formed as the Schwann cell encircles the axon
Vertebral
Intervertebral
foramen
body
Ventral
Cl-::>"_--.,
C2
C3
C4
Cervical
nerves
B
Vertebral
arch
Dorsal
Ventral
Thoracic nerves
Lumbar
plexus
Spinal nerve
Cauda
equina
~~
~~
MOlor
_ -->"",,--,-,'00'
I---~'-~~
Dorsal root
ganglion
Coccygeal
nerve
A
A, Schematic drawing of the spinal cord and the spinal
nerves (posterior view). The spinal nerves emerge from the
spinal canal through the intervertebral foramina. There are 8
pairs of cervi<al nerves, 12 pairs of thoracic nerves,S pairs of
lumbar nerves, 5 pairs of sacral nerves, and 1 pair of coccygeal nerves. Except in the region of the 2nd to the 11 th
thoraci< vertebrae (T2-T11), the nerves form complex networks called plexuses after exiting the intervertebral forClmina. Only the main branch of each nerve, the ventral ramus,
Sensory
rOOI
o
is depicted. Adapted from rorrord, GJ & Anllgnostako5. N.fl.
(J 984). Principles oj Anatomy and PhySiology (.1lh ed.J. New York:
Harper & Row. B, Cross-section of the cervi<al spine showing
the spinal cord in the spinal canal and the nerve roots exiting
through the intervertebral foramina. C, Cross-section of the
lumbar spine shOWing the nerve roots of the cauda equina in
the spinal canal. 0, Each exiting nerve root complex in the intervertebral foramen consists of a motor root, a sensory root,
and a dorsal root ganglion.
•
\vinds around it many times, pushing its cytoplasm
and nucleus to the outside layer. Unmyelinated gaps
called nodes of Ranvier lie between the segments of
the myelin sheath at approximately I to 2 I11Ill apart.
The myelin sheath increases the speed of the conduction of nerve impulses, and insulates and maintains the aXOIl. Impulses arc propagated along the
unmyelinated nerve fibers in a slow, continuous way,
whereas in the myelinaled nerve fibers the impulses
"jump" al a higher speed from one node of Ranvicr
to the next in a process called saltatOl)" conduction.
The conduction velocity of a myelinated nerve is directly proportional to the diameter of the fiber,
which lIsually ranges [Tom 2 La 20 J-l111. MOlar f-ibers
-. f:
innervate skeletal muscle have large diameters,
do sensor~y' fibers that relay impulses associated
touch, pressure, heat, cold, and kinesthetic
such as skeletal muscle tension and joint poSensory fibers that conduct impulses for dull,
diffuse pain (as opposed to sharp, immediate pain)
the smallest diameters. Nerve fibers are packed
closely in fascicles, \vhieh are further arranged into
bUIl1(II·es that make lip the nen'c itself. The fascicles
arc the functional subunits of the nCI-ve.
the same nerve. \!Vhere the nerves lie close to b
INTRANEURAL CONNECTIVE TISSUE OF
PERIPHERAL NERVES
Successive layers of connective tissue surround the
nerve fibers-called the endoneurium, perineuriUl11,
,'mel epineurium-and protect the fibers' continuity
(Fig. SA). The protective function of these connec-
Sensory nerve root
Sensory cell
body in dorsal
rool ganglion
!
Spinal nerve \
\
Motor cell body
in anterior horn
of spinal cord
Dorsal
\
Dorsal ./
ramus /
Ventral
Peripheral nerve ... '-
Motor nerve rool
I
nerve
fiber
1
~'-----------
Schematic representation of the arrangement of a typical
spinal nerve as it emerges from its dorsal and ventral nerve
roots. The peripheral nerve begins after the dorsal ramus
branches off. (For the sake of simplicity, the nerve is not
shown entering a plexus.) Spinal nerves and most peripheral nerves are mixed nerves: they contain both sensory
(afferent) and motor (efferent) nerve fibers. The cell body
and its nerve fibers make up the neuron. The cell bodies of
the motor neurons are located in the anterior horn of the
spinal cord, and those of the sensory neurons are found in
the dorsal root ganglia. Here, a motor nerve fiber is shown
innervating muscle and a sensory nerve fiber is depicted
innervating skin. Adapted from Rydevik, B., Brown, MD., &
Lundborg, G. (/98·4). Pathoanatomy and pathophysiology of
nerve root compression.
?
tive tissue la)rers is essential becausc nerve flbcrs
extremely susceptible to stretching and comp
sion.
The outcnTlOst layer, the epineurium, is loc
between the fascicles and superficially in the ne
This rather loose connective tissue layer serves
cushion during movements of the nel\'e, protec
the fascicles from external trauma and maintain
the oxygen supply sy'stem via the epineural bl
vessels. The amount of epineural connective ti
varies among nerves and at different levels wi
Spine, 9, 7.
.... ,':;.-
..
or pass joints, the epineurium is often more ab
c1ant than e1scwhere, as the need for protection
be greater in these locations. The spinal nerve r
are devoid of both epineuriurn and perincuri
and the ncrvc fibcrs in the nel\'e root may therc
bc morc susceptible to trauma (Rydevik et aI., 19
The perineurium is a lamellar sheath that
compasses each fascicle. This sheath has great
chanical strength as well as a specific biochem
barrier. Its strength is demonstrated by the fact
the fascicles can be inflated by fluid to a pressur
approximately! 1000 mm of mercury (I-Ig) beFore
perineuI'ium ruptuI·es.
The barrier function or the perineurium che
cally isolates the l1el\'e fibers From their surrou
ings, thus preserving an ionic environment of th
terior of the fascicles, a special milieu interieur.
endoneurium, the connective tissuc inside the f
cles, is composed principally of fibroblasts and
lagen.
The interstitial tissue pressure in the fasci
the endoneurial fluid pressure, is normally slig
elevated (+ 1.5 ± 0.7 mm Hg [Myers & Pow
1981]) comparecl \\'ith the pressure in surround
tissues sLIch as subcutaneous tissue (-4.7 ± 0.8
Hg) and muscle tissue (~2 ± 2 mm Hg). The elev
endoneurial fluid pressure is illustrated by the
nomenon whereby' incision of the perineurium
sults in herniation of nerve fibers. The endoneu
fluid pressure may increase further as a resu
trauma to the nerve, with subsequent edema. S
a pressure increase ma.y affect the microcircula
and the function of the nerve.
THE MICROVASCULAR SYSTEM OF
PERIPHERAL NERVES
The peripheral nerve is a well-vascularized struc
containing vascular networks in the epineurium
perineurium, and the endoneurium. Because
inlpulse propagation and axonal transport dep
Node of Ranvier
Schwann cell
Myelin
sheath
\
\
Axon
Node of Ranvier
Schematic drawings of the structural features of a myelinated nerve fiber. Ad(1fJted from
Sunderland 5. 0978j. Nerves and Nerve Injuries (2nd ed.). Edinburgh: Churchill Livings/one.
•
on a local oxygen supply, it is natural that the
microvascular system has a large reserve capacity.
The blood supply to the peripheral nerve as a
whole is provided by large vessels that approach the
nCl\'C segmentally along its course. \,Vhen these local nutrient vessels reach the nerve, they divide into
ascending and descending branches. These vessels
run longitudinally and frequently anastomose with
the vessels in the perineurium and endoncuriul11.
Within the epineurium, large arterioles and venules.
50 to 100 fl-l11 in dial11etel~ constitute a longitudinal
vascular system (Fig. SA).
\Vithin each fascicle lies a longitudinally oriented
capillary plexus with loop formations at various levcis. The capillary system is Fed by arterioles 25 to
150 fLrn in diameter that penetrate the perineurial
membrane. These vessels run an oblique course
through the perineurium, and it is believed that because of this structural peculiarity, they arc easily
closed like valves in the event that tissuc prcssurc
inside the fascicles increases (Lundborg, 1975;
Myers et aI., 1986). This phenomenon may explain
why even a limited increase in endoncurial fluid
pressure is associated with a reduction in intrafascicular blood flow.
The built-in safety system of longitudinal anastomoses provides a wide margin of safety if the regional segmental vessels arc transectcd. In an ex-
Epineurium
Perineurium
Endoneurium
Schematic drawing of a segment of a peripheral nerve.
The individual nerve fibers are located within the en·
doneurium. They are closely packed in fascicles, each of
which is surrounded by a strong sheath, the perineurium.
A bundle of fascicles is embedded in a loose connective tissue, the epineurium. Blood vessels are present in all layer
of the nerve. A, arterioles (shaded); V. venules (unshaded).
The arrows indicated the direction of blood flow. Adapted
from Dahlin, L.B., Rydevik, B., & Lundborg, G. (1986). The
parhophysiology of nerve entrapments and nerve compression
injuries. In A.R. Hargens (Ed.). Effects of lv1echanical Stress on
Tissue Viability. New York: Springer-Verlag.
•
peri mental animal in vivo model, it is extremely difficuh to induce complete ischemia to a nerve by 10surgical procedures. For example, if the whole
sciatic··til."'1i nerve complex of a rabbit (15 em long)
is surgically Sepal"aled from its surrounding structures and the region"11 nutrient vessels arc ClIt. there
is no detectable reduction in the intrafascicular
blood flow as studied by intravital microscopic tcch·
Y"iGucs. Even if such a mobilized nerve is ellt distally
or proximall~'. the intraneural longiwdinal vascular
systems cnn maintain the microcirculation at least 7
l~ 8 ern rTom the clil end. If a non mobilized nerve is
0"'.,'" cut, there is still perfect microcirculation even at the
vcr")' tip of the nerve; this phenomenon demonstrates the sufficiency of the intraneural vascular
collalcrals. Howevcr~ othcr sllldies in rats indicate
.that stripping the epineural circulation rronl nerve
bundles causes demyelination of subpel'ineural
nerve fibers.
"swdling" of the most caudal part of the respective
dorsal neryc root, called the dorsal rOOl ganglion.
The dorsal root ganglia arc locatcd in or close to the
intervertebral Foramen. Unlike the nCn!c roots, the
dorsal root ganglia arc not enclosed by cerc
brospinal fluid and the meninges. Instead. they are
enclosed by both a multilayc:rcd connective tissue
sheath, similar lo the perineurium of the peripheral
w
4
2
Anatomy and Physiology of
Spinal Nerve Roots
~
t
I
I
f
!
3
In the earl~' embryological developmental stages,
the spinal cord has the same length as the spinal
column. However, in the full~: grown individual, the
spinal cord ends as the COIlUS mcdullaris, approximately at the level of the fit'st lumbar vertebra. A
nerve root that leaves the spinal canal through an
intervertebral foramen in lhe lumbar or sacral spine
therefore has to pass from the point where it leaves
the spinal corel. which is in the lower thoracic spine.
lO the point of exit li'om the spine (Fig. 5-5). Because the spinal cord is not present be.low the first
lumbar vertebra, dte nervous content of the spinal
canal is only comprised of the lumbosacral nerve
roots. This "bundle" of nerve roots within the IUI11bar and sacral pan of the spinal canal has been suggested to resemble the tail of a horse and is therefore
often called the cauda equina, that is, lail of horse.
Two different lypes of nerve roots are ("ound
within the lumbosacral spine, ventral/motor roots
and dorsal/sensory roots. The cell bodies of (he motor axons are located in the anterior horns of the
gray malleI' in the spinal cord, and because these
nerve roots leave the spinal cord from the ventral aspect, they are also called ventral roots. The other
type of nelVC root is the SenSOl)', or dorsal root: As
the name suggests, these nerve roots mainly COll1
prise sensory (Le., afferent) axons and reach lhe
spinal cord at the dorsal region of the spinal cord.
The cell bodies of the sensory axons are located in a
2
..
3
7
5
8
6
The intraspinal nervous structures as seen from behind.
The vertebral arches are removed by cutting the pedicles
(1). A ventral (2) and a dorsal (3) nerve root leave the
spinal cord as small rootlets (4). Before leaving the spinal
canal, the dorsal root forms a swelling called the dorsal
root ganglion (5), which contains the sensory cell bodies.
before forming the spinal nerve (6) together with the
ventral nerve root. The nerve roots are covered by a cen·
tral dural sac (7) or with extensions of this sac called
nerve root sleeves (B). ReprodtlCec/ with permission ;,om
O/marker. K. (J 99/ J. Spinal nerve rOOI compression. ACt/Ie
compression of the cauda equina sludied in pigs. Acta Orthop
w
Scand. 62. Suppf 242.
•
nerve, and a loose connective tissue layer called
epineurium.
""Vhen the ncnlC root approaches the intervertebral foramen. the root sleeve gradually encloses the
nerve tissue more lighLly. The subarachnoid space
and the amount of cerebrospinal fluid surrounding
each nen'c root pair will thus become gradually reduced in the caudal direction. Compression injury
of a nerve root may induce an increase in the permeability of the endoneurial capillaries. resulting in
edema formation (Olmarker et aI., 1989b; Rydevik &
Lundborg, 1977), This can lead to an increase of the
imraneural fluid and subsequent impairment of the
nutritional transport to the nerve (Myers & Powell.
1981; Myers, 1998), Such a mechanism might be
particularly important at locations where the nerve
roots are tightly enclosed by connective tissue. Thus
there is a more pronounced risk for an "entrapment
syndrornc" within the nerve roots at the intcrvenebral foramen than more central in the cauda equina
(Rydevik et aI., 1984), The dorsal root ganglion, with
its content of sensory nCt've cell bodies, tighlly enclosed by meninges, might be panicularly susceptible to edema formation.
MICROSCOPIC ANATOMY OF SPINAL NERVE
ROOTS
There are two microscopically different rc.:::gions of
the ncnlC roots. Closest to the spinal cord is a central glial segment comprised of glial cells and therefore resembles the microscopic organization of central nervous structures at the spinal cord or the
brain. This glial segment is transferred to a nonglial
segment in a "dome-shaped" junction a few millimeters from the spinal cord. This nonglial segment
is organized in the same manner as the endoneurium of the peripheral nerves, that is, with
Schwann cells instead of glia cells. However, some
small islets of glia cells also are found in this otherwise "peripherally" organized endoncurium.
MEMBRANOUS COVERINGS OF SPINAL NERVE
ROOTS
The axons in the endoneurium are separated from
the cerebrospinal fluid by a thin layer of connective
tissue called the root sheath, This root sheath is the
structural analogue to the pia mater that covers the
spinal cord. There are usually 2 to 5 cellular layers
in the root sheath, but as many as 12 layers have
been identified, The cells of the proximal part of the
outer layers of the root sheath arc similar to the pia
cells of the spinal cord, and the cells in the distal
part are more similar to the arachnoid cells of the
spinal dura. The inner layers of the rOOl sheath are
comprised of cells that show similarities to the cells
of the perineurium of peripheral nerves. An inlerrupted basement membrane encloses these
cells separately. The inner layers of the root sheath
constitute a diffusion barrier bet\veen the cl1do~
neurium of the nerve roots and the cerebrospinal
fluid, This barrier is considered to be relatively
weak and may only prevent the passage or macromolecules.
The spinal dura encloses the nerve roots and the
cerebrospinal fluid, When the two layers of the cranial dura enter the spinal canal. the outcr layer
blends with the periosteum of the pan of the laminae of the cervical vertebrae facing the spinal canal.
The inner layers join the arachnoid and become the
spinal dura. [n contrast to the rooL sheath, the
spinal dura is an effective difTusion barrier: The barrier properties are located in a connective tissue
sheath between the dura and the arachnoid called
the neurotheliunl. Sirnilar to the inner layer or the
root sheath, this ncurothclium resembles the
perineurium of the pel'iphcral nen'es. It is suggested
thal these t\VO layers in facl form the perincuriulll
when the nerve rooL is transformed to a peripheral
nerve upon leaving the spinal.
THE MICROVASCULAR SYSTEM OF SPINAL
NERVE ROOTS
Information about the vascular anatomy of the
nerve roots has mainly been derived from studies
on the vascularization of the spinal corel. Therefore.
the nomenclature of the various vessels has been
somewhat confusing. A SLlllllTIat·y of the existing
knowledge on nerve root vasculature will be presenteel below,
The segnlental arteries generally divide into three
branches when approaching the intervertebral foramen: (1) an anterior branch that supplies the posterior abdominal wall and lumbar plexus. (2) a posterior branch that supplies the paraspinal muscles
and facet joints, and (3) an intermediate branch that
supplies the contents of the spinal canal. A branch
of the intermediate branch joins the nCI"ve root at
the level of the dorsal root ganglion, There are usually three branches from this vessel: one La the ventcal root, one to the dorsal root, and one to the vasa
corona of the spinal corel.
:7:.
!I
The branches to the vasa corona of lhe spinal
corel, called medullal)' arteries, arc inconsistcnt.
7 to 8 remain or the 128 from the cmbryologiperiod of life, and each supplies more than one
segment or the spinal core!. The main medullary
in the thoracic region of the spine was discovered by Adamkiewicz in 1881 and still bears his
name. The medullary arteries run parallel to the
~~-nt:rve roots (Fig. 5-6). In humans, there are no connections between these vessels and the vascular net
work of the nerve roOlS. Because the medullmy
feeder artcrics only occasionally supply lhe nerve
roots with blood, they have been referred to as the
extrinsic vascular system of the cauda equina.
The vasculature of thc nerve roots is formed by
branches from the intermediate branch of the segmental arlclY distally and by branches from lhe
vasa corona of the spinal cord proximally. As opposed to the medullat)' arteries, this vascular network has been named the intrinsic vascular system
of the cauda equina. The dislal branch to the dorsal
root first forms lhc ganglionic plexus within the dorsal root ganglion. The vessels run within the outcr
layers of the root sheath, called cpi-pial tissuc. As
there are vesscls coming from both distal and proximal directions, the ncr've roots are supplied by two
separate vascular systems. The two syslems anastomose at approximalely two thirds of the nerve rool
lenglh from the spinal corel. This location demonstrates i:\ region of a less-developed vascular nClwork
and has been suggested lo be a panicularly vulnerable site of lhe nerve rools.
The arterics of the intrinsic system send branches
down to the decper parts of thc nerve tissue in a Tlike manner. To compensate for elongation of the
nerve roots, the m·teries are coiled both longitudinally and in thc steep running branches bClwccn the
different fascicles (Fig. 5-6). Unlike peripheral
nerves, the vcnules do not course togethcr with the
arteries in the nerve roots but instead usually have
a spiraling course in the deeper pans of the nerve.
There is a barrier of the endoneurial capillaries in
periphcral nerves called the blooel-nerve barrier,
which is similar to the blood-brain barrier of lhe
central nervous system (Lundborg, 1975; Rydevik &
Lundborg. 1977). The presence of a corresponding
barrier in netlle roots has been queslioned. If presem, a blood-nerve barrier in nerve rOOlS does not
seem to be as well developed as in endoneuria I capillaries of peripheral nerve, which implies that
edema may be formed more easily in nen.,'c roots
lhan in peripheral nelves (Rydevik el aI., 1984).
M
Schematic presentation of some anatomical features of t
intrinsic arteries of the spinal nerve roots. The arterioles
within the cauda equina may be referred to either the e
trinsic (1) or the intrinsic (2) vascular system. From the su
perficial intrinsic arterioles are branches that continue al
most at right angles down between the fascicles. These
vessels often run in a spiraling course, thus forming vasc
lar "coils" (3). When reaching a specific fascicle they
branch in a T-like manner, with one branch running cranially and one caudally, forming interfascicular arterioles
(2b). From these interfascicular arterioles are small
branches that enter the fascicles, where they supply the
endoneurial capillary networks (2c). Arterioles of the extrinsic vascular system run outside the spinal dura (4) and
have no connections with the intrinsic system by local va
cular branches. The superiicial intrinsic arterioles (2a) are
located within the root sheath (5). Reproduced with permi
sion from Olmarker. K. (1991 J. Spinal nerve rooe compression.
Acuee compression of the cauda equina seudied in pigs. Acta
Qnhop Scand, 52. Supp1242.
Biomechanical Behavior
Peripheral Nerves
of
External trauma to the extremities and nerve e
trapment may produce mechanical deformation
the peripheral nerves that resulls in the deterior
tion of nerve Function. If the mechanical trauma e
ceeds a certain degree, the nerves' buill-in mech
nisms of prolcction may not be sufficient, resulli
in changes in nerve sLructure and function. Co
. mon modes of nerve injlll)' are stretching and co
pression. which may be. inflicted, respectively,
rapid eXlension and crushing.
STRETCHING (TENSILE) INJURIES OF
PERIPHERAL NERVES
ties, and the nerve behaves 1110re !ike a plastic m
terial (I.e., its response to the release of loads is i
complete recovery).
Although variations exist in the tensile streng
of various nerves, the maximal elongation at t
elastic limit is approximately 20(~fi, and comple
structural failure seen1S to occur at a maximu
elongation of approximately 25 to 30(k;. These v
ues are for normal nerves; injury to a nell/e may
duce changes in its mechanical properties, name
increased stiffness and decreased elasticit\'.
Stretching, or tensile, injuries of peripher
nerves are usually' associated with severe acciden
such as when high-energ.\' tension is applied to t
brachial plexus in association with a birth~relat
injur.\'. as a result of high-speed vehicular co!lisio
or after a fall from a height. Such plexus injuri
may result in partial or total functional loss of som
or all of the nerves in the upper extremity, and t
consequent functional deficits represent a conside
able disability in terms of sensory and motor lo
The outcorne depends on which tissue componen
of the nerves are damaged as well as on the exte
of the tissue injury'. Of clinical importance is the o
servation that there can be considerable structur
damage (perineurial sheath injuries) induced
stretching with no visible injury' on the surface
the nerve (Case Study 5-1).
Nerves arc strong structures with considerable ten-
sile strength. The maximal load that
tained
b~/
call
be
SlIS-
the median and ulnar nerves is in the
rangc of 70 to 220 newtons (N) and 60 to 150 N, rcspectively. These f-igurcs afC of academic interest
only because severe intraneural tissue damage is
produced b:v tension long before a nerve breaks.
A discussion of the elasticity and biomcchanical
properties of nerves is complicated by the fact that
nerves are not homogeneous isotropic I11aterials but,
instead, composite structures, \vith each tissue component having its O\VI1 biomechanical properties.
The connective tissues of the epineurium and perineurium arc primarily longitudinal structures.
\,Vhen tension is applied to a nen'e, initial elongation of the nerve under a ver.v small load is followed by! an interval in \vhich stress and elongation
show a linear relationship characteristic of an elastic material (Fig. 5-7). As the limit of the linear re~
gion is approached, the nerve fibers start to rupture
inside the endoneurial tubes and inside the intact
perineurium. The perineurial sheaths rupture at approximately 25 to 30i?-c elongation (ultimate strain)
above in vivo length (Rydevik et aI., 1990). Aftcr this
point, there is a disintegration of the elastic proper-
"
(L
~
~
~
~
iii
12
11
10
9
8
In-situ strain: 11.0 :!:. 1.5%
In-situ stress: 0.05 :c: 0.3 MPa
7
6
5
4
3
2
1
0
Ultimate stress: 11.7 :.:: 7 MPa
Ultimate strain: 38.5 :::: 2.0'%
0
4
8
12
16
20
24
28
32
36
40
Strain ('%)
The stress-strain behavior of a rabbit tibial nerve. The nerve
exhibits a low stiffness toe region of approximately 15 %
and begins to retain significant tension as the strain increases
beyond 20%. Reproduced with permission from Rydevik. 8.L.,
K\·van, M.K., Myers, R.R., et al. (1990). An in. vitro mechanical an
histological study of acute stretching on rabbit tibial nerve. J Orthop Res, 8, 694-707.
"--
,-""
----0
I
. B~ichial Plexus Palsy
"ri.~.i-ipg
the birth process, a newborn suffered a trac-
.~fit9~. injury in his left brachial plexus. A few months
.-0'.>
.1,:
'~t~~:·.h~'presents with the upper left arm in a static posi'~i?r-.()fadduction, internal rotation of the shoulder, exi~h?i{)pof the elbow, pronation of the forearm, and flexi'9r'l9fthe wrist. He does not respond to sensory stimulus
'iri;~i~>shoulder and presents biceps and brachioradialis
~',-"~fje.x.ia. A sudden deformation and high tensile stress in·
. j.. '~'. 'f',""'·.
rr~~.in the (5-(6 nerve roots affected the mixed (motor
:'.'~·ilnd·.s.en.sory) neural functions, mainly the muscles respon·",-slbh~,'for the scapulohumeral rhythm (see Chapter 12).
.1' ;:\)'AHE.rb's palsy is diagnosed. The sudden elongation
J7;r~uHer~~d during the traction can lead to structural dam~;.~ge an9 reduction in the transverse fascicular cross-sec{ftio.f1~(area, producing impairment of the intraneural
:~",,:vascu.lar flow and impulse transmission.
1', ·.In less severe cases, functional restoration may occur
t.- within·weeks or months. In more severe cases, healing may
"i·i:~~X~k(place during the first 2 to 3 years, but if the structural
:~r~erveJnjury is severe, considerable long-term functional
:s<'disability can result, If structural derangement of the nerve
:i·::i.trunk has taken place, nerve grafting may be required.
~.
):i:' , "
J
High-energy plexus injuries represent an extreme
type of stretching lesion caused by sudden violent
trauma. A differcm stretching situation of considerable
clinical interest is the sutllling of the two ends of a CuL
nerve under moderate tension. This situation occurs
when a substantial gap exists in the continuity of a
nerve trunk and the restoration of the continuity requires the application of tension to bring the neryc ends
back togethel: The moderate, gradual tension applied
to the nerve in these cases may stretch and angulate local feeding vessels. It ma.y also be sufficient to reduce
the transverse fascicular cross-seclional area and impair the intraneural nutritive capillary flow (Fig. 5-8).
As the sutured nelve is stretched, the perineurium
l'ightens; as a result, the endoneurial fluid pressure is
increased and the intrafascicular capillaries may' be
obliterated. Also, the flow is impaired in lhe scgmcn·
lal, feeding, and draining vessels, as it is in larger
vessels in the epineudulll, and at a certain stage lhe
intraneural microcirculation ceases. Intravital observations of intraneural blood now in rabbit tibial
nel'ves (Lundborg & Rydevik, 1973) showed that an
elongation of 80M induced impaired venular no\\' and
that even greater tension produced continuous impairment of capillary and arteriolar now until. at
15% elongation, all intraneural microcirculation
ceased completely. For the same nerve, an elongation
(strain) of 60/0 induced a reduction of nerve action potential amplitude by 709'0 at I hour with recover)' to
normal values during I-hour restitution. At 12
elongation, conduction was completely blocked by
hour and showed minimal recovery (Wall et a
1992). Such data have clinical implication in ner
rcpail~ limb trauma, and limb lengthening.
A situation of even more gradual stretching, a
plied over a long time, is the growth or intraneu
tumors such as schwannomas. In this situation, t
nerve fibers are forced into a circumferential Cou
around the gradually expanding tumor. Function
changes in cases of sllch very gradual stretching a
oftcn minimal or nonexistent.
COMPRESSION INJURIES OF PERIPHERAL
NERVES
It hns long been known that compression of a ner
can induce symploms stich as numbness, pain, a
muscle weakness. Thc biological basis for the fun
tional changes has been investigatcd eXlensiv
(Rvdevik & Lundborg, 1977; Rydevik et aL, 1981),
these investigations (Fig. 5-9), even mild compre
sion was observed lo induce structural and fun
lion'll changes, and the significance of mechanic
Elongation
Schematic representation of a peripheral nerve and its
blood supply at three stages during stretching. Stage I: T
segmental blood vessels (S) are normally coiled to allow
for the physiological movements of the nerve. Stage II: U
der gradually increasing elongation, these regional vess
become stretched and the blood flow in them is impaire
Stage III: The cross-sectional area of the nerve (represen
within the circle) is reduced during stretching and the in
traneural blood flow is further impaired, Complete cessa
tion of all blood flow in the nerve usually occurs at approximately 15% elongation. Adapted from Lundborg. G.
Rydevik. 8. (1973). ENeas 0; strecching the tibial nerve of the
rabbit: A preliminary study 0; (he intraneural circulation and
barrier (uncrion of the perineurium. J Bone Joint Surg, 558. 3
Recording
electrodes
and amplifier
r1./,...,/
Stimulating
electrodes
Compression
chamber
I
Compressed
ajr inlets
Schematic drawing of an experimental setup for studying
deterioration of nerve function during compression.
Adapted from Dahlin, L.B.• Rydevik, 8.• & Lundborg. G. (1986).
The pathophysiology of nerve entrapments and nerve compres'
sian injuries. In A.R. Hargens (Ed.). Effeers of Mechanical Stress
on Tissue Viability. New York: Springer- Verlag.
sure levels (approximately 30 to 80 mm Hg) may in
ducc intrancural edema. \vhich in turn may becom
organized into a fibrotic scar in the ncn'e (Ryclevi
& Lundborg, 1977),
Compression at approxirnately 30 111m Hg als
brings about changes in the axonal transport sys
tems, and long-standing compression may thus lea
to depletion of axonally transported proteins dista
to the compression site. Such blockage of axona
transport induced by local compression (pinching
may cause the axons to be marc susceptible to ad
ditional compression distally. the so-called double
crush syndrome.
Slightly higher pressure (80 mm Hg, for example
causes complete cessation of intraneural blood flow
the nerve in the locally compressed segment be
comes completely ischemic. Yel, even after 2 hour
or more or compression, blood flo\\' is rapidly· re
stored whcn the pressure is released (Rydevik ct aI
1981), hen higher levels of pressure (200 to 40
mm Hg, for example) applied directly to a nen'c ca
induce structural nerve fiber damage and rapid de
terioration of nerve function, with incomplete rc
covel)· after even shortel· periods of compression
Hence, the magnitude of the applied pressure an
the severity of the induced compression lesion ap
pear to be correlated.
factors such as pressure level and mode of compression became apparent.
Mode of Pressure Application
Critical Pressure Levels
Experimental and clinical observations have re·
vealed some data on the critical pressure levels at
which disturbances occur in intraneural blood flow,
axonal transport, and nerve function. Certain pressure levels seem to be well defined with respect to
structural and functional changes induced in the
nerve, The duration of the compression also influences the development of these changes,
At 30 mm Hg of local compression, functional
changes may occur in the nerve, and its viability
may be jeopardized during prolonged compression
(4 to 6 hours) at this pressure level (Lundborg et aI.,
1982), Such changes appear to be caused by impairment of the blood flow in the compressed part
of the nerve (Rydevik et aI., 1981), Corresponding
pressure levels (approximately 32 111111 Hg) were
recorded close to the median nen.re in the carpal
tunnel in patients with carpal tunnel syndrome,
while in a group of control subjects the pressure in
the carpal tunnel averaged onl~,.. 2 mm Hg. Longstanding or intermittent compression at lo\\' pres-
The pressure level is not the only faclOr that influ
ences the severity of nerve injlll)' brought about b
compression, Experimental and clinical evidenc
indicates that the mode of pressure application
also of major signiflcance. "Its importance is illu
trated by the fact that direct compression of a nerv
at 400 mm Hg by means of a small inflatable cu
around the nCI've induces a more severe nerve injur
than does indirect compression of the nerve at 100
mrn Hg via a tourniquet applied around the extrem
ity. Even though the hydrostatic pressure acting o
the nerve in the former situation is less than ha
that in the latter, the nerve lesion is more sever
probably because direct compression causes a mor
pronounced deformation of the nerve (especially a
its edges) than does indirect compression, in whic
the tissue layers between the compression devic
and the nerve "bolster" the nerve, One nlay also con
clude that the nerve injllly caused by compression
not directly related to the high hycfrostatic pl'essu
in the center of the conipressed nerve segment bu
instead is more dependent on the specific mechan
cal deformation induced by the applied pressure,
Mechanical Aspects of Nerve Compression
;!:
I
Electron microscopic analysis of the deformation of
the nerve fibers in the peroneal nerve or the baboon
hind .limb induced b~,' tourniquet compression
demonstrated the so-called edge effect; that is. a
specific lesion was induced in the nerve fibers at
both edges of the compressed nerve segment: the
nodes of Ranvicr were displaced toward the noncompressed parts of the nct·ve. The nerve fibers in
the center of the compressed segment, where the
hydrostatic pressllt·c is highest. generally were nOl
affected acutely. Tht.: large-diameter nerve fibers
were usually affected, but the thinner fibers were
spared. This finding confirms theoretical calculations that indicate larger nerve fibers undergo a rdatively greater deformation than do thinner fibers at
a given pressure. It is also known clinically that a
compression lesion of a nCI-vc IIrst affects the large
fibers (e.g., those that can)' molOr function). while
the thin fibers (e.g.. those that mediate pain sensalion) arc often preserved. The intraneural blood vessels have also been shown to be injured at the cdges
of the compressed segment (Rydevik & Lundborg,
1977), Basically, the lesions of nerve libers and
blood vessels seem to be consequences of the pressllre gradient. which is maximal just al the edges of
the compressed segment.
In considering the mechanical effects on nCI-ve
compression, keep in mind that the effect of a given
pressure depends on the way in which it is applied,
ils magnitude and duration. Although pressure may
be applied with a variety of spatial distributions.
two basic types of pressure applications arc generally encountered in experimental settings and in
pathological conditions. One type is uniform pressure applied around the entire circumference of a
longitudinal segment or a nerve or exu'emity. This
is Ihe kind of purely radial pressure thai is applied
by the common pneumatic tourniquet. It has also
been used in miniature apparatlls to produce controlled compression of individual nerves (Rydevik
& Lundborg, 1977) (Fig, 5-9). Clinically, Ihis Iype of
loading on a nerve probably occurs when the pressure on the rnedian nenre is elevated in the carpal
tunnel. producing a characteristic syndrome.
Another type
mechanical action takes place
when the nerve is compressed laterally_ This is the
kind of defol-mation that occurs if a nelve or extremity is placed between two parallel nat rigid sur
faces that arc then moved toward each othel~
squeezing the nerve or extremity. This type of deformation occurs if a sudden blow by a rigid object
squeczes a nervc against the surface of an undcrly~
ing bone_ h may also occur when a spinal nerve
compressed by a herniated elisc (Case Study 5·2).
The details of the deformation or a nerve may
quite difTerent in these two cases of loading. In un
form circumferential compression like that appli
Sciatic Pain
35-year-old male construction worker has chronic
low back pain radiating below the- Ie-f! knee that is
more severe with lifting activities and prolonged posi,
tions. After a careful examination. certain neurological
signs were found. Positive straight leg raising and L5
motor and sensory functions were affected.
A MRI shows a herniated disc at level L4-L5 with
posterolateral protrusion. which laterally compresses (he
left L5 nerve root. Compression of !he nerve deforms it
toward a more elliptical shape. increasing strain and
stress loads_ The effects of the pressure and mechanical
deformation resultant from the load affects the nerve
tissue, its nutrition, and the transmission function. in·
flammation of the nerve root, induced by the nucleus
pu!posus, may sensitize the nerve rOOt so that mechanical ne-rve root deformation causes SCiatic pain (Case
A
Study Fig 5·2·1),
or
R
.:.--
~I
_.' Case Study Figure S-2·1.
y
factor at both high and low pressures, but ischemia
pl[t~·s a dominant role in longer-dunHion compression. This phenomcnon is illustraled b\' the fact that
direct ncr'vc compression at 30 111111 Fig for ) to 4
hours produces reversible changes, whereas pro·
longed compression a(}()\'c this time period at this
pressure level ma~' cause irreversible damage [0 the
nCI"e (Lundborg CI aI., 1982; Rydevik el aI., 1981),
Corn pression at 400 I11Ill Hg causes a much more severe nervc injur~1 arter 2 hours than after 15 minules. Such information indicalcs that Cn:n hiuh
pressure has to "acl" for a certain period of lime for
injur.v to occur. These data also give some informalion about the viscoelastic (time-dependent) properties of" peripheral nCI"\'t: tissue. Sufficient time OJ! f!51
elapse for permanent deformation to develop.
i
,
,
I
I
. .-:
,'.'
--I------1I-- X
....
-,'
";.:,,
• I' :'
..
,-",
"/':
",
" '; ,';""
'/,:,". :," '.: 'f;-i::;:
Pressure
dislribution ..--, ---c___
A
-
, .
---..."
"
'
.. "
:',:'
B'
C'
E
E'
F
i,'
G
. . '",.,.
I
F'
G'
x
,,
B
"iG, 5-11
A, Theoretical displacement field under lateral compression as a result of uniform clamping pressure. B, The origi,
nal and deformed cross-sections are shown for maximum
elongation in the x direction of 10. 30, and 50%. The vectors shown from A to fJ:, B to a', and so forth, indicate the
paths followed by the particular points A, B, and so forth
during the deformation.
I
•
or
that this kind
deformation can trigger r-iring of
nerves, resulting in a sensation or pain when the
ne"'e Gbers are laterallv compressed. The details or
such deformation of nerves and their functional consequences have not been studied extensively and require further research for their elucidation.
Duration of Pressure Versus Pressure Level
Knowledge is limited regarding the relalive importance o[ pressure and time, rcspectively, in the production of nerve compression lesions. Mechanical
faclors seenl to be rehllively more imporlant al
higher than at lo\ver pressures. Time is a significant
Biol17echal1ical Behavior (Jf
Spinal Nerve Roots
The nerve roots in the thee;:d sac lack epinei"fl-tl:lw
and perin~uril1m, but under tensile loading they exhibit both elasticity and tensile strength. The ultil11,tle load rnr ventral spinal nelye roots from the
thecal sac is belween 2 and 22 ! , and 1'01' dorsal
nerve roots from the thecal sac the load is between
5 and 33 N. The length of the nerve roots I'rorn the
spinal corel to the foramina \'aries From appro:dI11ntely 60 mm at the L I level to approxilmllely 170
mm at the S I level. The mechanical propertics of
human spinal nerve roots arc c1ifferelll for an~1 given
nerve roOl at its location in the cenlral spinal canal
ancl in the lateral intel'vertebral foramina. The ultinwtc load for the inrr4lthecal portion of human SI
nel'VC rOOIS at the S I level is appro:"\imatcl~; 13 N,
and thal 1'01' the roraminal portion is approximately
73 N. For human nCI'vc roots at the L5 level. the corresponding values arc 16 Nand 71 N, respectively
(Fig. 5-12). Thus, the "allles ror ultimate load are
approxil11«lel~" five times higher ror the foraminal
segment of the spinal ner'vc roots than for the intrathecal pan ion of the same nClVC roots under tensile loading. However, the cross-sectional area of the
nerve root in the intervertebral fOI'(llllcn is sie:nificantlv
larger than that of the same nerve root in lhe
.
thecal sac; thus, the ultimate tensile stress was more
comparable for the two locations. The ultimate
strain LInder tensile loading is 13 La 19% for the
human ne,Ye root at the L5-S I level (Fig. 5-13),
The nerve roots in the spine arc nol static slruclures: they 1110\'C relati\'e to the surrounding tissues
-
-
Ultimate load: Human sinal nerve roots
u
'" S
2
rn'"
§
5
o'--_-'-_-';:-=----'-'-_-'-_-';-;:"'-_-'-_..J
Sl
LS
N"4
N~4
Diagram illustrating values for ultimate load obtained for
human spinal nerve roots under tensile loading. INR, intrathecal nerve roots; FNR, foraminal nerve root. Note the
marked difference in ultimate load for the intrathecal and
the foraminal portions of the nerve roots. Error bars indi-
cate standard deviation. Reproduced with permission from
Weinstein, IN., Latviotte, R.. Rydevik, B., er al. (1989), Nerve. In
J, \IV, Frymoyer & S.L. Gordon (Eds.). NevI Perspectives on Low
Back Pain (Chapter 4, pp. 35-130). Park Ridge. IL: MOS. {Based
on
a workshop arranged by the National Institutes
roots normally' adhere lo the surrounding tiss
above and below the intervertebral disc they
n:rsc, compression may give rise to intraneural
sion, Spencer and associates (1984) measured
contact force between a simulated disc herniat
and a dclcwmed nerve 1'001 in cadavers. Taking
area of contact into aeCOUlll, they assumed a con
pressure of approximately 400 mm Hg. \<\'ith
duced disc height. the contact force and pressure
tween the experimental disc herniation and
nerve root was reduced, They suggested that th
findings may explain in pan why sciatic pain is
lieved after chcmol1ucleolysis, and as disc degene
tion progresses over time and the disc height ther
decreases,
[n central spinal stenosis, the mechanics or ne
root compression arc completely diFferent. Un
these conditions, the pressure is applied circum
cntially around the nerve roots in the cauda cqu
at a slow, gradual rate, These different deformat
factors, together with the fact that the nerve ro
centrally within the cauda equina differ comple
from the nerve roots located more laterally, clos
the discs, ma.y explain some of the dillcrcnt syrm
toms found in spinal stenosis and disc herniatio
of Hea/rh
Ultimate strain: Human spinal nerve roots
(NIH) in Airlie, Virginia, USA, May 1988.}
_ 1S
with every spinal motion. To allow for such motion
the nerve roots in the intervertebral foramina, for
example. must have the capacity to glide. Chronic
ilTitation with subsequent fibrosis around the nerve
roots. in association with conditions such as disc
herniation and/or Foraminal stenosis. can thus impair the gliding capacity of the nel"VC roots. This
produces repeated "microstretching" injuries of the
nerve roots even during normal spinal 1110Vcments,
which might be speculated to induce yet FUrl her tissue iITitation in the nerve root components, The
normal range of movements of nerve root.s in the
human lumbar spine has been measured in cadaver
experiments. It was found that straight leg raising
moved the nerve roots at the level of the intervertebral foramina approximately 2 to 5 mm.
Certain biomechanical Factors are obviollsly involved in the pathogenesis of val'ious symptoms induced by nen'e root deformation in association with
disc herniation and spinal stenosis and resulting in
radiating pain. In disc herniation, only one nerve
root is usually compressed, Because individual nerve
l
c
.~
en
10
r"n
§
5
S
o'----'---';:-J'-----''----'-------',-:-'---'-S1
LS
N.=4
N=4
Ultimate strain for human spinal nerve roots under ten
loading. INR, intrathecal nerve root; FNR, foraminal ne
root. Reproduced with permission from Weinstein, IN., L,1M
otte, R" Rydevik, 8., er at. (1989). Nerve. In I \tV Fryrnoyer &
Gordon (Eds.). New Perspectives on Low Bc1Ck Pain (Chapter
pp. 35-130). Park RieJge, Ii: MOS. IBased on cl workshop
arranged by the NatioTlallnstitut.es of Health (NIH) in Air/ie,
ginia, USA, May /988.J
COMPRESSION OF SPINAL
has been moderate interest in the past to
nerve root compression in experimental modEarly sludies in the 1950s and 1970s found that
roots seemed to be more susceptible La COI11-';l;'pl'essie,n than did peripheral nerves. During recent
howevcl~ the interest in nerve rool patho;ggph,ysiology has increased considerably and a oumof studies have been performed thal arc rc.vic,\Vc:ct below.
Some years ago. a model was presented to evalthe effects of compression of the cauda
in pigs, \\!hich for the first time allowed
experimental, graded compression or cauda
nCl"Ve rOOlS at known pressure levels (01marker. 1991) (Fig. 5-14). In this Illodel, ,he calida
equina was compressed by an inflatable balloon
that was fixed to the spine. The cauda cquina
"could also be observed through the translucent
balloon. This mndel made it possible to stud~' the
flow in the intrinsic nerve roOl blood vessels at
various pressure levels (Olmarkcr el aI., 1989a).
The experiment was designed in a way that the
pressure in the compression balloon \Vas increased by 5 mm Hg evcry 20 seconds. Blood flow
and vessel diameters of the intrinsic vessels could
simultaneously be obsClycd through the balloon
using a vilal microscope. The average occlusion
pressure for the arterioles was !"ound to be slightly
below and directly related to the systolic blood
pressure. and the blood Ilow in the capillary networks was intimately depcndent on the blood flow
of the adjacent venulcs. This corroborates the assumption that venular stasis may induce capillary
stasis and thus changes in the microcirculation of
the nerve tissue and is in accordance with previous studies in which such a mechanism has been
suggested as involved in carpal tunnel syndrome.
The mean occlusion pressures for the venules
demonstrated large val"iations. However, a pressure of 5 to 10 mOl Hg was found to be sufficient
for inducing venular occlusion. Because of retrograde stasis, it is not unlikely to assume that the
capillary blood flow will bc affected as well in slIch
situations.
In the same experimental SCt-lIp. the effects of
gradual decompression, after initial acute compression was maintained for only a short while. were
studied. The average pressure for starting the blood
lIow was slightly lower at decompression than at
comprcssion For artcrioles. capillaries. and vcnulcs.
_
. . . .'--->-->---
Schematic drawing of an experimental model. The cau
equina (A) is compressed by an inflatable balloon (8) t
is fixed to the spine by two l-shaped pins (C) and a ple
glass plate (0). Reproduced with permission from O/marke
Rydevik. 8., & Holm. S. (1989a). Edema (ormation in spina/
nerve roars inducecl by experimental. graded compression.
experimental srudy on the pig cauda eq'Jina ~'.Jirh special re
1
ence co differences in effects between rapid and slow onsec
compression. Spine. ld. 579.
>-------
Howevcc with this protocol a full restoration or
blood flow did not occur until thc compression
lowcred from 5 to 0 mill Hg. This observation
ther supports the previous hypothesis that vasc
impairment is present even at low pressure leve
A cornpression-induced impairmenl of the
culature may thus be one mechanism for nerve
dysfunction because the nutrition of the nerve
will be affected. However, the nerve rools will
derive a considerable nutritional supply via d
sion from the cerebrospinal nuicl. To assess
compression-induced eFfects on the tolal conlr
tion to the nerve rools. an experiment was desig
in which jH-labelcd melhyl-glucose was allowe
be transported to the nen/c tissue in the c
pressed segment via both the blood vessels and
cerebrospinal fluid diffusion aflcI- systemic in
tion. The results showed that no compensa
mechanism from cerebrospinal fluid diffusion
could be expeclr:d at the low pressure levels. On the
COil Lntl)·,
to
111m Hg compression was sufficient to
induce a 20 to 300/0 reduction of the transport or
mClhyl·glucosc to the nerve r001S, as compared
with the control.
\·Ve know 1"1'0111 experimental studies on peri phend nen!cs that compression Illay also induce an
increase in the vascular permeability, leading to an
intraneural edema formation. Such edema may' increase the cndoncuriul Fluid pressure, which in
turn may impair the endoneurial capillary blood
flow and jeopardize the nutrition of the ner'vc
rOOls. Because the edema Llsually persists for some
lime after the rClllo\'al of a compressive agenl.
edema may negativel~r affect the nerve root for n
longer period than the compression itself. The
presence of intraneural edema is also related to the
subsequent formation of intraneural fibrosis and
may therefore contribute to the slow recaVCI)" seen
in some patients with nellie compression disorders. To assess if' intraneural edema also may form
in nerve roots as the result of cornprcssion, the distribution of Evan's blue-labeled albumin in the
nerve tissue \vas analyzed after compression at various pressures and at various durations (Olmarker
el aI., 1989b). The study showed thal edema was
formed even at low pressure levels. The predomi·
nant location was at the edges of the compression
zone.
The function of the nelYC roots has been studied
by direct electrical stimulation and recordings either on the nenie itself or in the corresponding
muscular segments. During a 2- hour compression
period, a critical pressure level for inducing a reduction uf MAP-amplitude seems to be located between 50 and 75 111m Hg. Higher pressure levels
(100-200 mm Hg) may induce a tOlal conduction
block \vith var.ying degrees of recovery after compression release. To stud)! the effect's of compression on sensory nerve fibers, electrodes in the
sacrum were used to record a compound nerve action potential after stimulating the senso(v nerves
in the lail, that is, distal to the compression zone.
The resuhs showed thaI the sensory fibers are
slightly more susceptible to compression than are
the motor fibers. Also, the nerve roots are more
susceptible to compression injury if the blood
pressure is lowered pharmacologically. This further indicates the importance of the blood supply
to maintain the functional properties of the nerve
roots.
ONSET RATE OF COMPRESSION
One faclor that has not been fully recognized
compr-ession trauma of nerve tissue is the onset r
of the compression. The onset rate, that is. the ti
from start to full compression. may vnry clinica
rrom fractions of seconds in lraUI11atic condition
months or years in association with degenerat
processes. Even in the clinically' rapid-onset ra
there ma:v be a wide variation or onset rales. \V
the presented model. it was possible to vary the
set time of the applied compression, Two onset ra
have been investigated, Either the pressure is p
sent and compression is started by flipping
switch of the compressed-air system used to inl
the balloon or the compression pressure leve
slowl~1 increased during 20 seconds. The first on
rate was measured at 0.05 to 0.1 seconds. thus p
viding a rapid inflation of the balloon and a ra
compression onset.
Such a rapid-onset rate has been found to ind
more pronounced effects on eden1a formati
mcth~ll-glucose transport. and impulse propagat
lhan the slow-onser rale (Olmarker, 199 I). Rega
ing methyl-glucose transport, the results show t
the levels within the compression zone arc m
pronounced at the rapid than at the slow onSd r
at corresponding pressure levels. There was als
striking difference between the two onset ra
when considering the segments ourside the co
pression zones. 1n the slow-onset series, the le
approached baseline values closer to the compr
sion zone than in the rapid-onsel series. This m
indicate the presence of a more pronounced ed
zone edema in the rapid-onset sel"ics, with a sub
quent reduction of the nutritional transport in
nerve tissue adjacent to the compression zone.
For the rapid-onset compression, which is lik
lo be more closely related to spine trauma or d
herniation than to spinal stenosis, a pressure of
mill Hg maintained for only 1 second is surflcien
induce a gradual impairment of nc,'ve conduct
during the 2 hours studied aftcr the cornpress
was ended. Overall, the mechanisms for these p
nounced differences between the different on
rates arc not clear but may be related to dilTeren
in the displacement rates of the compressed ne
tissue toward the uncompressed parts, as a resul
the viscoelastic propcnics of the nellie tissue, S
phenomena may ICfld nOl only ro structural dam
to the nerve fibers but also to stn.lcturfll change
the blood vcssels with subsequcnt edema formati
The gradual formation or intraneural edema may
also be closely" related to observations of a graduall.\'
increasing dilTcrcncc in nerve conduction impair-
ment between the two onset rates (Olmar-ker et al..
1989b).
MULTIPLE LEVELS OF SPINAL NERVE ROOT
COMPRESSION
.'
.
••1.
!'atients with double or multiple levcls of spinal
stenosis seem to have morc pronounced symptoms
than do patients with stenosis only at one level. The
presented model was modified to address this interesting clinical question. Using two balloons at two
adjacent disc levels. which resulted in a 10-mm uncompressed nerve segment bctwcen the balloons,
induced a much more pronounced impairment of
nerve impulse conduction than previollsl~' had been
round at corresponding pressure levels (Olmarker &
Rydevik. 1992). For instancc. a pressurc of 10 mm
I-1g in two balloons induced a 60% reduction of
nerve impulse amplitude during 2 hours of COIllpression, whereas 50 111m I-Ig in one balloon showed
no reduction.
The mechanism ror the diFrerence between single
and double compression may not simply be based
on the fact thal the nerve impulses have lo pass
more than one compression zone at double-level
compression. There may also be a mechanism
based on the local vascular anatomy of the nerve
roots. Unlike for peripheral nerves, there are no re·
gional nutritive arteries from the surrounding
structures to the intraneural vascular system in
spinal nerve roots. Compression at two levels might
therefore induce a nutritionally impaired region betwcen the two compression sites. In this way. the
:.~: segment nffected by the compression would be
"', widened from one b,~lIoon diarneter (10 mm) to two
balloon diameters including the inLCrjacent nerve
segment (30 mm). This hypothesis was partly confirmed in an experiment on continuous analyses of
the total blood flow in the ul1compresscd nerve scg·
ment located bet ween two compression balloons
(Takahashi et al.. 1993). The results showed that a
64% reduction of total blood now in the uncoll1pressed segment was induced when both balloons
were inflated to 10 mm Hg. At a pressure close to
the systemic blood pressure lhere was complete ischemia in the nerve segment. Thus, experimental
evidence shows that the blood supply to the nerve
segment located between two compression sites in
nerve roots is severely impaired although this Tlel'VC
segment itself is ullcomprcssed. Regarding n
conduction, the effects were much enhanced i
distance between the compression balloons wa
creased frorn one \'cn~bral segment La two v
bral segments (Olmarker & Rydevik. 1992). Thi
dicates that the functional impairment may
directl:v related to the distance between the
compression sites.
CHRONIC NERVE ROOT COMPRESSION IN
EXPERIMENTAL MODELS
Th~
discLlssion of compression-induced effect
nerve roots has deall primarily wiLh acute c
pression, that is, compression that lasts for s
hours and with no survival of [he animal. To be
mimic various clinical situations, compres
must be applied over longer periods of time. T
arc probably man." changes in the nerve Li
such as adaptation ofaxons and vasculature,
will occur in patients but cannot be studied in
perimental models using only I to 6 hours of c
pression, Another important factor in this con
is the onset rale that was discllssed previousl)
clinical syndromes with nerve root compress
lhe onset time may in many cases be quite s
For instance, a gradual remodeling of the v
brae to induce a spinal stenosis probably lead
an onset lime of many years. Jt will of cours
dirficult to mimic such a situation in an exp
mental modcl. It will also be impossible to
control over the pressure acting on the nerve r
in chronic models because or the remodeling
adaptation or the nerve tissue to the applied p
sure. However, knowledge of the exact pressur
probably of less importance in chronic tha
acute compression situations. Instead, chr
models should induce a conLrolled compres
with a slow onset time that is easily reproduc
Such models may be well suited for studies
pathophysiological events as well as inlcrven
b~' surgery or drugs. Some attempts have b
made to induce such compression.
Dclamarlcr and collaborators (1990) present
model on dog cauda equina in which they appli
constricting plastic band. The band was tighte
nrollnd the thecal sac to induce a 25, 50, or 750/
duction of the cross-sectional area. The band
left in place for various limes. Analyses were
~forl1led and showed both stnlctural.and functi
changes that were proportional to the degree of
striction.
Experimental study to analyze the effects on nerve conduc·
from Cornefjord, 1''11., 5ato, K., O/marker, K., et 011. (1997),
tion velocity of nucleus pulposus (1), the combination of
nucleus pulposus and compression (2), and compression
only (3). The nucleus pulposus and the constrictor were applied to the first sacral nerve root in pigs. The contralateral
nerve root served as a control. Reproduced with permission
morJe! for chronic nerve root compression studies. Present
To induce a s!c)\ver onset and more controlled
compression, Cornefjord and collaborators (1997)
used a constrictor to compress the nerve roots in
the pig (Fig. 5-15). The constrictor was initially! intended for inducing vascular occlusion in experimental ischemic conditions in dogs. The constrictor consists of an outer metal' shell that on the
inside is covered with a material called amaroid
that expands when in contact with fluids. Because
of the metal shell, the amaroid expands inwards
with a maximum expansion after 2 weeks, resulting in compression of a nerve root placed in the
central opening of the constrictor. Compression of
the first sacral nerve root in the pig resulted in a
significant reduction of nerve conduction velocity
and axonal injuries using a constrictor with a defined original diameter. An increase in substance P
of a porcine mode! (or controlled slow-onset compression
analyses of anatomic aspects, compression onser rate, and
phologic and neurophysiologic effects. Spine, 22. 946-957
in the nen'(' root and the dorsal root ganglio
lowing such compression also has been fo
Substance P is a neurotransmitter that is re
to pain transmission. The study may thus pr
experimental evidence that compression of
roots produces pain.
The constrictor model has also been use
stud.\' blood flow changes in the nerve root v
lature. It could then be observed that the b
Flow is not reduced just outside the compre
zone but significantly reduced in parts o
nerve roots located inside the constrictor. In
context, note that in case of disc herniation
nerve root may become sensitized by substa
from the disc tissue (nucleus pulposus) so
mechanical root deformation can induce
nounced sciatic pain.
',:",f;1>Thc peripheral nerves arc composed of nerve
}t;,,:rs, layers or connective tissue, and blood vessels.
;~{2 The nCIYC f'ibers arc extremely susceptible 10
ruuma but because they arc surrounded by sliccessive layers of connective tissue (the epineuriulll and
1?~rinc·urillm)J they arc 111cchanically protected.
;:.'$§';}Strctching induces changes in intraneural
bl~~d noll' m~d nerve Iiber S[;'llctllre before the
trunk ruptures.
Comrression of a nerve can cause ItlJury to
nerve fibers and blood vessels in the nerve,
'mail,lv at the edges of the compressed nerve segV>:ment. but also by ischemic mechanisms.
j"."'j
5- Pressure level, duration of compression. and
mode of preSSllre application arc signillcant variables in the development of nen'c injury,
6.. Spinal nerve roots arc anatomically different
from peripheral nerves and therefore react differently to mechanical defonnalion,
T: Spinal nerve roots arc morc susceptible than
peripheral ncrves to mechanical dcformation,
mainly because of the lack of protective connective
tissue layers in nerve roots.
Myc:rs. R.R .. ~Itlrabmi, H., .& "(}WI:I!. H.C. (1986), Red
uel"\'l' hlood flow in l.'delllatous nCllropathics-A bi
chanica I lIle(,:h.lOism . .11;LTol'tlscu!al' Ues. 32. 145-151
Myi.'rs, R.R" 8.: Powell, H.e. (1981). Endoncuri,d fluid
sure in peripheral neuropathies. In A.R. Hargens
Ti..; SIIC flilid PreSSl/re tlml C01lllJOsithm (p. 193). Balti
Willi.II11S &. Wilkins.
Olillarkl"', K., Rydc\·ik. B., &: Holm. S. (1989a). Edema r(
lion in spin£llnl'l'\'l' roots induced by experimelltal. g
compression, An experirnenud Sllld~' on lhe: pig c
equina with special reference 10 dilTerences in effec
1\\,(.'(.'11 rapid .. Iud slow onset of compression. Spiw:, I..J
Olmarkcr. K., Rydc\'ik. B.. Holm. S .. 1..'1 ill. (l9&9b). EffcC
CXpl'l'illlental graded compression on blood flow in s
nl,.·l'vc roms. r\ vilal microscopic study 011 Ihe po
caudn equinll. J Orthop Res, 7,8Ii.
Olmarkcl'. K. {1991l. Spinal nerve rOOl compression.
compression of dle cauda equina studied in pigs . .. IeU
f/mp Scallll, 62. Suppl 2·n.
Ollllarker. K. 6.: Rydc\·ik. B. (1992). Single: \'('rstls doubk·
COllllll·l,.·ssion. '\n experimental slud.... un till..' porcine c
l'qUill,1 with 'lll.t1ySl'S of nen·c impulse condu<:lion pr
lies. Cfill On/lOp. 279, 3539.
Olmarker. K. & Hasuc, :\1. (1995). Classificalion and p
physiology of spinal pain syndromes. III J.1'$. \\\:inSl
B. Rydl'\'ik (Eds.), Essc/llial...
Ihe Spillt'. RJ\'l'n
!\l,.·w York. NY.
R~·(!l'\'ik. B.L., Kw~n, \I.K., ;-"'Iyel's, R.R .. l'l al. (1990).
vitro mechanical and histological study of <ll..:lHe stret
on r;,hbit tibial nerve . .I Onhop Ucs. B, 694-701.
Ryd(·\·ik. B. & Lundborg. G. (1977). Pl:rIllt.'ahilily of
ncur:ll micl'U\'l'ssels :lnd perineurium following :
gradl,.·d cxpl'l'illlenlal ncrn: cOlllprcs~ion. SC(/Iu} J Na:>
or
COI1Slr Sl/".~.
REFERENCES
CorndjonJ. lVi., Saw. K.. Olmarkcr. K.. ci :d. (1997). A nHldcl
ror chronic n('IT":" roOI compression studies. PrCSi..'IlI~llion
of a porcine model for contrnllt·d slow-onset comprl'ssion
\\"ith an.tlyscs of anatomic ::lSPCCIS. comprl,.·ssion onSel
r~lle, and morphologic and neurophysiologic crfeels.
SpilIC, 22. 946-957,
[hillin. L.B .. Ry(kvik, B.. &: Lundborg. G. (1986). Thc p.IIIHI'
physiology of nelTe eHtrapmellis and nervi..' cOlllpr":"ssion
injuries. In A.R. Hargens (Ed.). E!l;'cIS
.\kclwllic:al
Stress 011 Tisslle Viability. New York: Springer-Verlag.
Delamartcr, R.B" Bohlman. I-U-I., Dodge, L.D" el a!. (1990).
Experimental lumbar spinal SlCllosis. Analysis of the cortic~d evoked potentials, microvasculature and histop<1thCII·
ogy. J BOIlt: JOillf SI/r:.:, 72:\. 110-120.
Lundbor!!, G., &: Rvdcvik, B. (1973), Effecls of strctchin!! Ihe
libial ~\cl"\"e of lill,.· rabbic: :\ prdimin.ary slUdy of tilt: i;llraneural cin:ul~lliol1 and Ihe b.uTicr funclion of Ihe per·
inelll'iul11. J BOllI: JOill1 SlIrg, 558, 390.
Lundborg. G. (1975). Structure :\Jld flllll:tioll of the inlra·
ncural micrO\'essl,:ls as related lO Iraum:l. edcma formation
and nerve fUllclion.) !JOIII.' )oim Surg. 57..1. 938.
Lundborg, G., e{ .11, (1982). Median nel'\'e compression ill the
C,I1'p,ll tunnel: The fUllclion:l1 respons(.· 10 cxpcrillH.'nlally
induced controlll·d pressure. .f f-/mu} SlIrg, 7, 152.
Myers, R.R. (1998). :\Iorphology or the peripheral nervous
System nnd its relntiollship to neuropathic pain. In T.L
';aksh. C. Lynch III. W.M. Zapol.:-"-1. ~'1azc, J.E Bicbuyck . .&
L.J. Saidman (Eds."J. Aw:...tht'S;u: Bio}o~ic FouI/datiolls (pp.
483-514). Phil'lCldphia: LippiIKOll·R~I\,l,.'n.
or
II. 179.
Rydc\'ik, B., Lundborg. G., &. Bagge, U. (1981). Eff(,.,
gra<k·d compr('ssion nn inlr,1I1cural blood flow. An in
sllld.... 011 rabbil tihi.tl ner\'e. J flawl Sure;. 6. 3.
Ryd":"\'ik, B.. Bro\\'ll, :\I.D .. &: Lundborg, G. (1
P<llhoanatolll~' and pathoph ....siolog.\· of nel'\'e root
pression. Sp;Ilt.'. 9. 7.
Ryd(.·\'ik. B.L.. Kw<tn, M.K .. Myers R.R., et :11. (1990).
vitro mechanical :\n(\ histological study of ,H:llle strel
on r .. hbil tibi:ll nen·l'. J Orrhop U,·.... 8. 694.
Spencer. D.L.. ~'lillcr. J.I\., .& BCriolini. J.E. (198·.J). Th
fects of intervertebral disc space narrO\\"ing nil the co
force bl·twcen the Ill'rn' rOOt and a simlll:llcd disc p
sion. Spille 9, 422.
SUl1ckrbnd. S. (1978). Nen'c.>: alld Nave III;tlri(~s (2nd
Edinburgh: Churchill Livingslone.
Takah;lshi, K.. Olm:-.rkc.·r. K., Holm, S .. cl £II. (1993). Do
I('vel cauda ('quina cqmprcssion. An experimenlal
wilh conlinuous monitoring of intraneural blood fl
Onhop Res, II, 104,
TOrlora, GJ., & Anagnoslakos. N.r. (1984). Prillci"l
AlUlIOI1lY (/ltd Physio}ogl' (41h cd.). N(.·w York: Harper &
Wall. E.J .. M:ISSic, J.B.. Kwan. M.K .. CI 011. (1992). E
lllcnta! stretch neurop'lthy. Changcs in nerve condu
under tension,) BOllt' .foitlf Stl,.~, N-B. 126,
Weinstein. J.N" L.li\lotll'. R.. Rydcdk, B. el 011., (1989). N
In J.W. Frymoyer 8.: S.L. Gordon (Eds.)_ Nell' 1'('I"spccti
1.0\1' Back Paill (Chapll'r 4, pp, 3S-130). Park Ridg(
'\AOS. [Based 011 a workshop ,IIT,lngcu by Ihe N:llion
stiwtes of Health (NIH) in AirJil'_ Virginia, USA. May
Nerve
structure
Structural
composition
and function
affected
Protects
from
external
trauma
-Biochemical
barrier
properties
'Preserves ionic
environment
'Axon's
protector
-Nutrition
Delivery
of blood
..
Local
oxygen supply
D
Illness or injury
_
Peripheral nerve's structure and alteration. Clinical examples.*
"This flow (hart is designed for classroom or group discussion. Flow chart is not meant to be exhaustive.
Returned
blood flow
Tl<lnsmission of
impulses;
intl<lcellular
tl<lnspon of
material
Of importJ.nce
(or impulse
transmission
Myelin
production and
synthesis of
trophic factors
Biomechanics o
Skeletal Muscle
Tobias Lorenz, Marco Campel/o adapred fr
Mark I. Pieman, Lars Peters
Introduction
Composition and Structure of Skeletal Muscle
Structure and Organization of Muscle
Molecular Basis of Muscle Contraction
The MOtor Unit
The Musculotendinous Unit
Mechanics of Muscle Contraction
Summation and Tetanic Contraction
Types of Muscle Contraction
Force Production in Muscle
Length-Tension Relationship
Load-Velocity Relationship
Force·Time ReJ(ltionship
Effect of Skeletal lvIuscie Architecture
Effect of Prestretching
Effect of Temperature
Effect of Fatigue
Muscle Fiber Differentiation
Muscle Injuries
Muscle Remodeling
Effects of Disuse and Immobilization
Effects of Physical Training
Summary
References
Flow Charts
IJ~S'.r'
Introduction
The muscular sYSlCll1 consists of three III usc Ie.:: l.ypes:
the cardiac muscle, which composes the hc[trt; the
Sl1100th (nonstriated or involuntary) muscle, which
lines the hollow internal organs; and the skeletal
(striated or voluntary) muscle, which auaches to the
skeleton via the tendons. The focus of this chapter is
the role and function of skeletal muscle.
Skeletal muscle is the most abundant tissue in
the human body, accounting for 40 to 45% of the tolal body weighL. The human body has morc than
430 skeletal !TIuscles, found in pairs on the right and
left sides of the body. The most vigorolls movements
are produced by fewer than 80 pairs. The muscles
provide strength and protection to the skeleton by
distributing loads and absorbing shock; they enable
Ihe bones to move at the joints and provide the
maintenance or body posture against force. Such
abilities usuall~' represent the action of muscle
groups, not of individurll mllscles.
The skeletal muscles perform both dynamic and
static work. Dynamic work permits locomotion and
the positioning of the body segments in space. Static work maintains body posture or position. In this
chapter we describe the composition and structure
of skeletal muscle. the mechanics of muscle contraction. force production in muscle. Illuscle fiber
differentiation, and muscle remade-ling.
Composition and Structure
o( Skeletal Muscle
An understanding of the biorncchanics of Illuscle
function requires knowledge of the gross anatomical structure and function of the musculotendinous
unit and the basic microscopic structure and chemical composition of the muscle fiber.
STRUCTURE AND ORGANIZATION
OF MUSClE
The slructuralunit of skeletal muscle is thc muscle
fiber, a long cylindrical cell with many hundreds of
nuclei. Muscle fibers range in thickness from approximately 10 to 100 fim and in Icngth from approximately 1 to 30 cm. A muscle fiber consists of
many myofibrils, which are invested by a delicate
plasma membl·ane called the sarcolemma. The sarcolemma is connected via vinculin- and dystrophinrich costameres with the sarcon1ctric Z lines, which
represent a pan of the cxtramyofibrillar cytoskcle-
ton. The myofibril is made lip of several sarcome
that contain thin (actin), thick (myosin), clas
(titin), and inelastic (ncbulin) filaments. Actin a
myosin are the contractile part of the myofibr
whereas titin and nebulin nrc part of the intramy
fibrillar cytoskelcton (Stromer et aI., 1998). T
myof-ibrils are the basic unit of contraction.
Each fiber is encompassed b:y a loose connect
tissue called the endomysium and the fibers arc
ganized into various-sized bundles, or Fascicles (F
6·1, A & B), which arc in (urn encased in a den
connective tissue sheath known as the perimysiu
The muscle is composed of several fascicles s
rounded by a fascia of fibrous connective tis
called the epimysium.
In general, each end of a muscle is attached
bone by tendons, which have no active contract
properties. The muscles form the contractile co
ponent and the tendons the series clastic comp
nent. The collagen fibers in the perimysium a
epimysium are continuous with those in the t
dons: together these fibers act as a stnlctural fr,\I
work for the attachment of bones and muscle fibe
The perim)'sium, endomysium, epimysium, and s
colemma act as parallel elastic components. T
forces produced by the contracting muscles
transmitted to bone through these connective
sucs and tendons (Kassel'. 1996).
Each muscle fiber is composed of a lal'ge nu
ber of delicate strands, thc mvol'ibrils. These
the contractile clements of muscle. Their structu
and function have been studied cxhausti\'c1y
Iighl and eleclron microscopy, and [heir his
chemistr~' and biochclllistry have been explain
elsewhel'c (Alvidson et aI., 1984; Guvton, 198
Approximately I p.m in c1iamctcl: the myofibrils
parallel to each other within the cytoplasm (s
coplasm) of the muscle fiber and extend throug
out its length. They vary in number from a few
several thousand depending on the diameter of
ll1uscle fibel~ which depcnds in turn on thc type
ll1usclc fiber.
The transverse banding pattern in striated m
cles repeats itself along (he length of (he mus
fiber. each repeat being known as a sal'comerc (F
6-IC). These striations are caused by thc individ
myofibrils, which are aligned continuously throu
Ollt the muscle fiber. The sarcomere is the fu
tional unit of the contractile syslem in muscle. a
the events that take place in one sarcomere arc
plicated in the others. Various sarcomere build
myofibril, various myof"ibrils build the muscle fi
and various muscle fibers build the muscle.
1
Schematic drawings of the structural organization of muscle. A, A fibrous connective tissue fascia. the epimysium.
surrounds the muscle, which is composed of many bun-
dles. or fascicles. The fascicles are encased in a dense con·
/
nective tissue sheath, the perimysium. B, The fascicles are
composed of muscle fibers, which are long, cylindrical,
multinucleated cells. Between the individual muscle fibers
are capillary blood vessels. Each muscle fiber is sur·
rounded by a loose connective tissue called the endomysium. Just beneath the endomysium lies the sarcolemma.
Epimysium
a thin elastic sheath with infoldings that invaginate the
fiber interior. Each muscle fiber is composed of numerous
delicate strands-myofibrils, the contractile elements of
muscle. C. Myofibrils consist of smaller filaments that form
a repeating banding pattern along the length of the myofibril. One unit of this serially repeating pattern is called
a sarcomere. The sarcomere is the functional unit of the
contractile system of muscle. D, The banding pattern of
the sarcomere is formed by the organization of thick and
thin filaments, composed of the proteins myosin and
actin, respectively. The actin filaments are attached at one
end but are free along their length to interdigitate with
the myosin filaments. The thick filaments are arranged in
a hexagonal fashion. A cross-section through the area of
overlap shows the thick filaments surrounded by six
equally spaced thin filaments. E, The lollipop-shaped mol·
ecules of each myosin filament are arranged so that the
long tails form a sheaf with the heads, or cross-bridges,
projecting from it. The cross· bridges point in one direction along half of the filament and in the other direction
along the other half. Only a portion of one half of a filament is shown here. The cross-bridges are an essential element in the mechanism of muscle contraction, extending
outward to interdigitate with receptor sites on the actin
filaments. Each actin filament is a double helix, appearing
as two strands of beads spiraling around each other. Two
additional proteins, tropomyosin and troponin, are associated with the actin helix and play an important role in
regulating the interdigitation of the actin and myosin filaments. Tropomyosin is a long polypeptide chain that lies
in the grooves between the helices of actin. Troponin is a
globular molecule attached at regular intervals to the
tropomysin. Adapted from Williams, P & War".vick. R, (1980).
Gray's Anatomy (36ch ed., pp. 506-515). Edinburgh: Cht/rdlill
A
S',glc [
muscle
Ilber
(cetl)
B
c
Thin
Cross section a
level of a band
IllamentlJl
(actin)
.,,''''':''
,
ThICK
o
filament
(mYOSin)
.;....
: " ,,':'.'
", ,
',::
.
:.:: :
' .....
Sa(comer,:, organilalion
Myosin {
Mamenl
E
\
Troponin
Livingstone.
Each sarcomere is composed of the Following:
I, The thin filaments (approximately 5 nm in di-
ameter) composed of the protein actin
2, The thick filaments (approximately 15 nm in
diameter) conlposed of the protein myosin
(Fig, 6-1, D & E)
3, The elastic filaments composed of the protei
titin (Fig, 6-2)
A, The inelastic filaments composedof the proteins nebulin and tilin
Aetin, the chief component or the thin filamen
has the shape of a double helix and appears as tw
strands of beads spiraling around t.:ach othcl: Two
additional proteins, troponin and tropomyosin, are
important constituents of the actin helix because
thev appear to regulate the making and breaking of
cOI~taclS between the actin and myosin filamel1ls
during contraction. Tropomyosin is a long polypeptide chain that lies in the grooves between the heiices of actin. Troponin is a globular molecule atlaehed at regular intcr'vals to the tropomyosin (Fig.
6-1, D & E).
The thick filaments are located in the central region of the sarcomere, where their ordcrl.y, parallcl
arrangemcl1l gives rise to dark bands known as A
bands because they arc strongly anisotropic. Thc
thin filaments arc atlachcd at either end of the sarcomere to a structurL' known as the Z line, which
consists of shon elements that link the thin filaments of adjacent sarcomcres. defining the limits
of each sarcomere. The thin filaments cxtcnd From
the Z line toward the center of the sarcomere,
where they overlap with the thick fibmcnts. Recently it was shown thal there is a third set or n1YOfibril filaments in the vertebrate striated Illuscles.
This connecting filament, named titin, links the
thick Filaments with the Z line (elastic I band region of titin) and is part of the thick filaments (A
band region o[ titin). This filament maintains the
central position of the A band throughout contraction and relaxation and might act as a template
during myosin ~,sselllbly.
Myosin, the thicker filament, is composed of individual molecules, each of which resembles a lollipop
with a globular "head" projecting from a long sha
or "tail." Several hundred slIch molecules are pack
tail to tail in a sheaf with their heads pointed in on
direction along half of the filament and in the opp
site dircction along the other half, leaving a hca
free region (the H zone) in between. The globul
heads spiral about the myosin filament in the regio
where actin and myosin ovcrlap (the A band) and e
tcnd as cross-bridges to interdigitate with sites o
the actin filaments, thus forming the structural an
I'unctionallink between the two filament types.
The intramyofibrilh.,r cvtoskeleton includes i
elastic nebulin filaments, which span from the
line to the actin filaments. Ncbulin might also act
a template for the thin filament assembly.
Titin is I j.1m long. It is the largest polypepti
and spans from the Z line to the IVI line. Titin is a
clastic filament. The part between the Z linc an
myosin has a string-likL' appearance. Titin has be
suggested to contdbutc greatly to the passh'c for
development of muscle during stretch (Fig. 6·2).
also might act as a template for the thick filame
assembly (Linke et aI., 1998; Squire et al., 199
Stromer et aI., 1998),
The I bane! is bisected by the Z lines, which conta
the portion of the thin filaments that does not overl
with the thick filaments and the clastic part of tit in.
the center of the A ballet. in the gap between the en
or the thin filaments, is the H zone, a light band co
taining only thick filaments and that part of tilin th
is integrated in the thick filaments. A narrow, da
area in the center of the H zone is the M line, 1'1
M-line
A-band
part
01 titin
Extent of one
!!lin molecule
The arrangement of titin molecules within the sarcomere. Adapted from Craig. R.(J994). The
stftJ(ture of the contract filaments. In A.G. Engel & Fraflzini·Armscrong (eds.). !vlyology (2nd ed.•
p. 150). New York: McGr<)w-Hill, Inc.
•..-,r; •.. __.
allel to the myofibrils and lend to enlarge and f
at the level of the junctions between the A an
bands, forming transverse sacs, or the terminal
ternae, that surround the individual myofi
completely,
The terminal cisternae enclose a smaller tub
tha1 is scpanllcd from them by its own membra
The smaller tubule and the terminal cisternae ab
and below it are known as a triad. The enclo
duced by transversely and longitudinally oriented
proteins that link adjacent thick filaments, maintaining their parallel Hrrangclllel1L The various areas of
the banding patlern arc apparent in the photomicrograph of human skeletal muscle shown in Figure 6-3.
Closely correlated with the rcpealing pattern or
the sarcomcrcs is an organized network of tubules
and sacs known as the sarcoplasmic reticulum.
The tubules of the sarcoplasmic reticulum lie par-
A
----"'....
Sarcomere
r
_--~""
_~
} Sarcoplasmic
reticulum
Mitochondrion {
I~M
--- Line
t
z
Line
H
Zone
B
A. Single muscle fiber with three protruding myofibrils. B. Electron photomicrograph of a cross·section of human skeletal m
c1e. The sarcomeres are apparent along the myofibrils. Characteristic regions of the sarcomere ar~ indicated.
Myo,libril
Mitochondrion
Sarcolemma
'l
Band
~
ill
E
0
u
m
A
Band
If)
I
Band
I z·
Triad
I
!
\j
Sarcotubule Terminal
cisternae
DD--------
1
Diagram of a portion of a skeletal muscle fiber illustrating
the sarcoplasmic reticulum that surrounds each myofibril.
The various regions of the sarcomere are indicated on the
left myofibril to show the correlation of these regions with
the sarcoplasmic reticulum, shown surrounding the middle
and right myofibrils. The transverse tubules represent an
infolding of the sarcolemma, the plasma membrane that
encompasses the entire muscle fiber. Two transverse
tubules supply each sarcomere at the level of the junctions
of the A band and I bands. Terminal cisternae are located
on each side of the transverse tubule, and together these
structures constitute a triad. The terminal cisternae connect with a longitudinal network of sarcotubules spanning
the region of the A band. Adapted from Ham, A.W & Cormack, D.H, (7979). Histology (8th ed.J. PiJifadefphia: IS, Lippincott
tubule is part or the transverse tubule system, or T
system, which are invaginatio!1s of the surface
membrane of the fiber. This membrane, the sarcolemma, is a plasma membrane that invests every
striated muscle (Fig. 6-4).
Molecular Basis
Contraction
of Muscle
The most widely..' held theory of muscle contraction
is the sliding filament theory, proposed simultaneously by A.F. Huxley and H.E. Huxley in 1964 and
subsequently refined (Huxley, 1974). According to
this theory, active shortening of the sarcomere, and
hence of the muscle, results from the relative m
ment of the actin and myosin filaments past on
other while each retains its original length.
force of contraction is developed by the m
heads, or cross-bridges, in the region of overla
tween actin and myosin (the A band), These c
bridges swivel in an arc around their fixed pos
on the surface of the m,vosin filament, much
the oars of a boat. This movement of the c
bridges in contact with the actin filaments
duces the sliding of the actin filaments tc)\var
center of the sarcomere, A muscle fiber con
\vhen all sarcomere shorten simultaneously
all-or-nothing fashion, which is called a twitch
Because a single movement of a cross-bridge
duces only' a small displacement of the actin
ment relative to the myosin filament, each in
ual cross-bridge detaches itself from one rec
site on the actin filament and reattaches itself
other site further along, repeating the process f
six tirnes, "with an action similar to a man p
on a rope hand over hand" (Wilkie, 1968).
cross-bridges do not act in a s)-'nchronized ma
each acts independently, Thus, at any given mo
(mly approximately half of the cross-bridges ac
generate force and displacement, and when
detach, others take up the task so that shorten
maintained, The shortening is reflected in th
comere as a decrease in the I band and a decre
the H zone as the Z lines move closer togethe
width of the A band remains constant.
A key..' to the sliding mechanism is the calciu
(Ca~'), which turns the contractile activity on
off. Muscle contraction is iniliated \vhen calci
made available to the contractile elements
ceases when calcium is removed, The mecha
that regulate the availability of calcium ions
contractile machinery are coupled to electric e
occurring in the muscle membrane (sarcolem
An action potential in the sarcolemma provid
electric signal for the initiation of contractile
it,v. The mechanism by which the electric signa
gel'S the chemical events of contraction is kno
excitationNcontraction coupling.
\Vhen the motor neuron stimulates the m
at the neuromuscular junction (Fig, 6-5A) an
propagated action potential depolarizes the
cle cell membrane (sarcolemma), there is a
\vard spread of the action potential along
system. (Details of this process arc given in F
6-5, A-C and in Box 6-1, which summarize
events during the excitation, contraction, an
laxation of muscle, Figure 6-5D shows the
Molor
end
plate
A
B
(1 )
Synaptic
vesicles
(3)
o
c
Schematic representation of the innervation of muscle fibers.
A, An axon of a motor neuron (originating from the cell body
in the anterior horn of the spinal cord) branches near its end
to innervate several skeletal muscle fibers, forming a neuromuscular junction with each fiber. The region of the muscle
membrane (sarcolemma) lying directly under the terminal
branches of the axon has special properties and is known as
the motor end plate, or motor end plate membrane. The rectangular area is shown in detail in B. B, The fine terminal
branches of the nerve (axon terminals), devoid of myelin
sheaths, lie in grooves on the sarcolemma. The rectangular
area in this section is shown in detail in C. C. Ultrastructu
the junction of an axon terminal and the sarcolemma. Th
vagination of the sarcolemma forms the synaptic trough
which the axon terminal protrudes. The invaginated sarcolemma has many folds, or subneural clefts, which grea
crease its surface area. Acetylcholine is stored in synaptic
cles in the axon terminal. Band C. adapted from Brobeck, I
(Ed.) (/979). Best and Taylor's Physiological Basis of Medical Pra
(10(h ed., pp. 59-1/3). Baltimore: Williams & Wilkins. D. Cros
bridge cycle of muscle contraction.
Events During Excitation, Contraction, and Relaxation of Muscle Fiber
1. An action potential is initiated and propagated in a motor axon.
2. This action potential causes the release of acetylcholine
irom the axon terminals at the neuromuscular junction.
3. Acetylcholine is bound to receptor sites on the motor
end plate membrane.
4. Acetylcholine increases the permeability of the motor
end plate to sodium and potassium ions. producing an
1 1. Actin activates the myosin ATPase found on the myosin
cfOss-bridge, enabling ATP to be split (hydrolyzed.) This
process releaSES energy used to produce movement 01
the myosin
cross~br;dges:
A, I'll ' ATP --; A ' Iv1 + ADP + P,
12. Oar-like movements of the cross bridges produce rela4
tive sliding of the thick and thin filaments past each
other.
end-plate potential.
S. The end-plate potential depolarizes the muscle mem-
13. Fresh AT? binds to the myosin cross-bridge, breaking
brane (sarcolemma), generating a muscle action pOlen-
the actin-myosin bond and allowing the cross-bridge to
tiallhat is propagated over the membrane suriaee.
dissociate from aGin:
6. Acetylcholine is rapidly destroyed by
acetylcholinesterclse
A . rv1 .;-. ATP -) A + tvl . ATP
on the end plate membrclne.
7. The muscle aGion potential depolarize') the transverse
tubules.
8. Depolarization of the transverse tubules leads to the re-
14. The ATPase hydrolyzes the myosin ATP complex to rhe
M . AlP complex, which represents rhe relaxed state of
the sarcomere:
lease of calcium Ions from the terminal cisternae of the
M' ATP -) Iv1 ' ATP
sarcoplasmic reticulum surrounding the myofibrils.
These ions are released imo the sarcoplasm in the direct
vicinitY of the regulatory proteins tropomyosin and troponin.
9. Calcium ions bind to troponin, allowing movement of
the tropomyosin molecule away from the myosin receptor sites on the actin filament that it had been blocking
and releasing the inhibition tilat had prevented actin
from combining with myosin.
10. Actin (A) combines with myosin ATP (M-ATP). In this
15. Cycles of binding and unbinding of actin with the
myosin cross-bridges at successive sites along the actin
filament (steps 11, 12, 13, CH'\d 14) continue as long as
ii)e concentration of calcium remains high enough to
inhibit the action of the troponin-rropomyosin system.
16. Concentration of calcium ions falls as they are pumped
into the terminal cisternae of the sarcoplasmic reticulum
by an energy-requiring process that splits ATP,
17. Calcium diSSOCIates from troponin. restoring the in-
state, ATP has been hydrolyzed to ADP and phosphate
hibitory action of troponin-tropomyosin. The actin fila-
bUt the products are still attached to myosin (receptor
ment slides back and rhe muscle lengthens. In the pr~s-,
sites on the myosin cross-bridges bind to receptor sites
ence of ATp, actin and myosin remain in tll~ dissociated;
on the actin chain):
relaxed stale.
A + Iv1 ' ATP --; A ' M ' ATP
(,l.,!odiiied from Luciano er ,l!. (I 978j, In HUIl1"n FuncliOn <1nd Structure (Fig 5,5D/ New York: ,\,'1cGril~·/·Hi,ll: and
,jdiJpted from (I'dlg. R. 099.:1). Myology (2nd ed.. p. 162) New York: McGra:'1·;·/iIl.)
lural features between actin and the cross-bridges
of mvosin.)
THE MOTOR UNIT
The functional unit of skeletal 111uscle is the mOlor
unit. which includes a single motor neuron and all
of the muscle rlbers innervated by it. This unit is the
smallest part of the muscle that can be made to CO!l4
tract independently. \Vhcn stimulated, all mu
fibers in the mOlor unit respond as one. The fi
of a Illotor unit arc said to show an all-or·none
sponse to stimulation: the),..' contract either m
mally or not at aIL
The number of muscle fibers forming a -m
unit is closely I"elated to the degree of contro
quirccl of the muscle, In small muscles thal perf
vcr~' fine movements, such as the extraocular I
des, each motor unit may contain less than a dozen
muscle fibers; in large n'lUscies that perform coarse
movements. such as the gastrocnemius, the motor
unil may contain 1,000 to 2,000 muscle fibers.
The fibers or each motor unit arc not contiguous
but dispersed throughout the muscle with fibers of
other units. Thus. if a single motor unit is stimulatex!. a large ponion or the muscle appears to con·
(rae£. If additional motor units of the nerve inner·
vating the muscle are stimulated, the muscle
corHracls with greater force. The calling in of adeli·
tional motor units in response to greater stimulation of the motor nClvc is called recruitment.
THE MUSCULOTENDINOUS UNIT
The tendons and the connective tissues in and
around the muscle belly are viscoelastic structures
that help determine the mechanical characteristics
of whole muscle during contraction and passive ex·
tension. Hill (1970) showed that the tendons represent a spring-like elastic component located in se·
rics with the contractile component (the contractile
proteins or the m~yofibril, actin, and myosin), while
the epimysium, perimysium, endomysium, and sm'·
colcmma represent a second elastic component 10·
cated in parallel with the contractile component
(Fig. 6-6).
When the parallel and series elastic components
stretch during active contraction 01- passive extension of a muscle, tension is produced and energy is
stored; when they recoil with muscle relaxation, this
energy is released_ The series elastic fibers are more
important in the production of tension than are lhe
parallel elastic fibers (Wilkie, 1956). Several investigators have suggested that the cross-bridges of the
myosin filaments have a spring-like property and
also contdbutc to the elastic properties of muscle
(Hill, 1968).
The distensibility! and elasticity of the clastic com·
ponents arc valuable to the muscle in severnl wavs:
I. Thev tend to keep the muscle in readiness for
contraction and assure that muscle tension is
produced and transillittcd smoothly during
contraction.
2. They assure that the contractile elements return to their original (resting) positions when
contraction is terminated.
3. They may help prevent the passive overstretch of the contractile elements when these
elements are relaxed, thereby lessening the
danger of Illuscle injUl)'.
PEC
,OOOOOH
SEC
CC
I
I
The musculotendinous unit may be depicted as consisti
of a contractile component ((C) in parallel with an elas
component (PEQ and in series with another elastic com
nent {SEQ_ The contractile component is represented b
the contractile proteins of the myofibril, actin, and myo
(The myosin cross-bridges may also exhibit some elastic
The parallel elastic component comprises the connectiv
tissue surrounding the muscle fibers (the epimysium, pe
mysium, and endomysium) and the sarcolemma. The se
elastic component is represented by the tendons. Adap
hom Keefe. CA.. Neil, E., & Joels. N (1982). I.;lvsc!e and the
lle!Voll$syslem. In Samson 1/</rI911(5 f.\ppli(~cJ PhySiology (i3th
ed, fJP, 248~259). Oxford: Oxford Unlversiry Press
4. The viscous property of the series and para
lel clastic components allows lIk'lll to abso
energy proportional to the rate of force app
cation and to dissipate energy in a timc~
dependent manner. (For a discussion of vis
coelasticity, see Chaptcl- 4.)
Thb viscoLIs property. combined with the ela
properties of the musculotendinous unit. is dem
slI-atcd in everyday activities. For example, whe
person altempts to stretch and touch the toes,
stretch is initially elastic. As the stretch is held, h
ever, further elongation or the muscle results fr
the viscosity of the Illuscle-tendon structure,
the fingers slo\\'I~' reach closer to the noor.
Mechanics of iVluscle
Contraction
_Electromyography' provides a mechanism for ev
ating and comparing neural effects on muscle
the contractile activity or the muscle itself in v
and in vitro. Much has been learned by using c
tromyography Lo study various aspects or the contractile process, particularly the time relationship
between the onset of electrical activity in the muscle
and aClltal contraction of the muscle or muscle
fiber. The following sections discuss the mechanical
response of a muscle to electrical (neural) stimulation and the various ways in which the muscle contracts to move a joint, control its motion, or maintain its position.
time of the rnuscle so that little or no relax
can occur before the next contraction is ini
(Fig. 6-8).
The considerable gradation of COlllraction e
ited by wholc muscles is achic\·ed by the differ
IIL,L,U--.
SUMMATION AND TETANIC
CONTRACTION
f
'
The mechanical response of a muscle to a single
stimulus of its molar nerve is known as a twitch,
which is the fundamental unit of recordable muscle activity. Following stimulation there is an interval of a few milliseconds known as the latency
period before the tension in the muscle fibers begins to rise. This period represents the timc re·
quit'cd for the "slack" in the elastic components to
be taken up. The time from the start of tension development to peak tension is the contraction time,
and the time from peak tension ulltil the tension
drops to zero is the relaxation time. The contraction time and relaxation time vary an10ng muscles,
depending largely on the muscle fibcr makeup (de·
scribed below). Some muscle fibers contracl with
a speed of only 10 mscc; others ma~; take 100 Illsec
or longer.
An action potential lasts only approximately I to
2 msec. This is n small fraction of the time taken for
the subsequent mechanical response, or twitch,
even in muscles that contract quickly; thus it is possible for a series of action potentials to be initiated
before the first twitch is completed if the activity of
the motor axon is maintained. \'\Then mechanical responses to successive stimuli arc added to an initial
response, the result is knc)\vn as summation (Fig.
6-7). If a second stimulus occurs during the latencyl
period of the first muscle twitch, it produces no additional response and the muscle is said to be completely refractory.
The frequency of stimulation is variable and is
modulated by individual motor units. The greater
the frequency of stimulation of the muscle fibers.
the greater the tension produced in the muscle as
a whole. However, a maximal frequency will be
reached beyond which the tension of the muscle no
longer increases. \"'hen this maximallension is sustained as a result or summation, the muscle is said
to contract tetanically. In this case, the rapidity of
stimulation olltstrips the contraction~relaxation
o
t
tOO
200
300
400
500
t
I;+-<- - - I >
s\
82
S3
A
Ilh t~
0
100
200
300
400
I
500
I-tS3
t
S,
S2
B
!I~ ,~
0
t
S,
100
200
300
400
I
500
..tI.5,S3
Time (msec)
C
Summation of contractions in a muscle held at a cons
length. A, An initial stimulus (51) is applied to the mu
and the resulting twitch lasts 150 msec. The second (5
third (5 J ) stimuli are applied to the muscle after 200-m
intervals when the muscle has relaxed completely, thu
summation occurs. 8.5 1 is applied 60 msec after 51, w
the mechanical response from 5.. is beginning to decr
The resulting peak tension is greater than that of the
gle twitch. C. The interval between 5.. and 5J is furthe
duced to 10 msec. The resulting peak tension is even
greater than in S, and the increase in tension produce
smooth curve. The mechanical response evoked by 5J
pears as a continuation of that evoked .by 51' Adapted
Luciano, 0.5., Vander. AJ.• & Sherman. J.H. (1978). Huma
tion and Structure (pp. J J3-(36). New York: McGraw,Hill.
3
c
0
'iii
c
2
<l>
>--
<l>
>
'g
0;
a:
0
is
100
ts
Is
is
200
300
400
is
is
500
600
Jumnmrnmmnnrrmnnm
ssssssssssssssssssssssssssssssssss
700
800
900
1000
Time (msec)
Generation of muscle tetanus. As the frequency of stimulation (5) increases (Le., the intervals shorten from 200 to 100
msec), the muscle tension rises as a result of summation.
When the frequency is increased to lOO/second, summation
activity' of their motor units, in both stimulation frequency' and the number of units activated. The
repetitive twitching of all recruited motor units of a
muscle in an as:vnchronous manner results in brief
summations or more prolonged subtetanic or
tetanic contractions of the muscle as a whole and is
a principal Factor responsible for the smooth movements produced by the skeletal muscles.
TYPES OF MUSCLE CONTRACTION
During contraction, the force exerted by a contracting muscle on the bony lever(s) to which it is attached is known as the muscle tension, and the external force exerted on the muscle is known as the
resistance, or load. As the muscle exerts its force, it
generates a turning erFect, or moment (torque), on
the involved joint because the line of application of
the muscle force usually lies at a distance from the
center of motion of the joint. The moment is calculated as the product of the muscle force and the per~
pendicular distance between its point of application
and the center of motion (this distance is known as
the lever arm, or moment arm, of the force),
Muscle contractions and the resulting muscle
work can be classified according to the relationship
bet\veen either the muscle tension and the resistance to be overcome or the muscle moment generated and the resistance to be overcome, as shown in
Box 6-2 (Krocmcr e! 'II., 1990),
becomes maximal and the muscle contracts tetanically,
ing sustained peak tension. Adapled from Luciano, OS.
AI, 8 Sherman, IN (978). Human Function and Structure
I i 3-136). New York: McGraw-Hi//.
Although no motion is accomplished and n
chanical \vork is performed during an isom
contraction, muscle work (physiological wo
performed: energy is expended and is Illostly
paled as heat, which is also called the isornetric
production. All dynamic contractions involve
ma.y be considered an initial static (isometric)
as the muscle first develops tension equal to th
it is expected to overcome.
The tension in a muscle varies with the t.y
contraction. Isometric contractions pro
greater tension than do conccntric contrac
Studies suggest that the tension developed
eccentric contractIon may even exceed that
oped during an isometric contraction. Thes
ferences arc thought to be due in large part
vat'ying amounts of supplemental tension
duced in the series clastic component of the
cle and to diFferences in contraction time
longer contraction time of the isometric an
centric contractions allows greater cross~
formation by the contractile components, thu
mitting greater tension to be generated (
1987). More time is also available for this te
to be transmitted to the series elastic comp
as the muscle-tendon unit is stretched. The l
contraction time allo\vs the recruitment of
tional nlotor units.
Komi (1986) has pointed oul that concentri
metric, and eccentric muscle contractions se
_
Types of Muscle Work and Contraction
Dynamic work: Mechanical work is performed and joint mo-
4. Iseinenial (iso, constant; inertial, resistance) contraction:
tion is produced through the following forms of muscle con-
This is a type of dynamic muscle work wherein the resis
traction:
1. Concentric (con, together; cemrum, center) contrac-
constant. If the moment (torque) produced by the musc
tion: When muscles develop sufficient tension to over-
tance against which the muscle must contract remains
is equal to or less than the resistance to be overcome,
come the resistance of the body segment, the mus-
the mllscle length remains unchanged and the muscle
cles shorten and cause joint movement. The net
contracts isomelfically. If the moment is greater than th
moment generated by the muscle is in the same di·
resistance, the muscle shortens (contracts concentrically
rection as the change in joint angle. An example of a
and causes acceleration of the body part. Isoinertial con~
concentric contraction is the action of the quadriceps
traction occurs, for example, when a constant external
load is lifted. At the extremes of motion, the inertia of
the load must be overcome; the involved muscles contract isometrically and muscle torque is maximal. In the
midrange of the motion, with the inenia overcome, the
mU'jcles contract concentrically and the torque is submaximal.
5. Isotonic (iso, constant; tonic, iorce) contraction: This ter
is commonly used to define muscle contraction in which
the tension is constant throughout a range of joint m()~
tion. This term does not take into aCCOLlnt the leverage'"
eilects at the joint. Ho\,vever, because the muscle force
moment arm changes throughout the range oi joint mo
lion, the muscle tension must also change. Thus, isoton
muscle contraction in the truest sense does not exist in
the production of joint motion (Kroll, 1987).
Static work: No mechanical work is performed and posture
or joint position is maintained through the following form
muscle contraction:
1. Isometric (iso, constant; metric, length) contra(ti()~,;,. :~~s.
c1es are not always directly involved in the prod~qi~~:-Pr.
joint movements. They may exercise either a rest'ra(ning
or a holding action, such as that needed t~ ~aint~i~ th
body in an upright position in opposing the force of
gravity. In this case the muscle attempts to shorten (i.e.,
the myofibrils shorten and in doing so stretch the series
elastic componenr. thereby producing tension), but it
does not overcome the load and cause movement; instead. it produces a moment that supports the load in a
iixed position (e.g., maintains posture) because no
,"
change takes place in the distance between the mlJsc!le
points of attachment.
in extending the knee when ascending stairs.
2. Eccentric (e.g., out of-, centrum, center) contraction:
When a muscle cannot develop sufficient tension and
is overcome by the external load, it progressively
lengthens instead of shortening. The net muscle moment is in the opposite direclion irom the change in
joint angle. One purpose of eccenlric contraction is to
decelerate the motion oi a joint. For example, when
one descends stairs, the quadriceps works eccentrically to decelerate flexion of the knee, thus decelerating the limb. The tension that il applies is less than
the force of gravity pulling the body downward, but it
is sufficient to allow controlled lowering of the body.
3. Isokinetic (iso, constant; kinetic, motion) contraction:
This is a type of dynamic muscle work in which movement of the joint is kept at a constan I velocity. and
hence rhe velocity of shortening or lengthening of the
muscle is constant. Because velocity is held constant,
muscle energy cannot be dissipated through acceleralion of the body part and is entirely converted to a resisting moment. The muscle force varies with changes
in its lever arm throughout the range of joint motion
(Hislop & Perrine. 1967). The muscle contracts concentrically and eccentrically with different directions of
joint motion. For example, the i1exor muscles of a
joint contract concentrically during flexion and eccentrically during extension, acting as decelerators during
the latter.
occur alone in normal human movement. Instead,
one type of contraction or load is preceded by a different type. An example is the eccentric loading
prior to the concentric contraction that occurs at
the ankle frolll miclstance to toe-off during gait.
Because muscles normally' shorten or lengthen at
"at)!ing velocities and with valying amoullls of tcnsion,
perfolll1ance and measurement of isokinetic work require the lise of an isokinetic dynamon1ctcl: This device
provides constant velocity of joint motion and maximum cxtcIllal resistance throughout the range of 010lion of the involved joint. thereby requiJing maximal
muscle torque. The use of the isokinctic dynamometer
provides a method of selective training and measurement, but physiological movement is not simulated.
c
.2 '.0
m
I
I
I
c
"
f-
"
.~
<i
~,
I
I
0.5
I
I
0;
I
a:
I
1.27
I
I
1.65 2.0
2.25
Sarcomere Length (JIm)
==i/i==
2.25-3.6 I'm c==:jr=z=J~\:~::\g;
::~;::~:::~\
z
20-225"m=9!
M
i':::":
:::!::~
A
Z
Ir==
<1.65"m==j~F=
Force Production in Muscle
The total force that a muscle can produce is influenced by its mechanical properties, which can be
described by examining the length-tension. loadvelocity, and force-time relationships of the muscle
and the skeletal muscle architecture. Other principal factors in force production arc muscle temperature, muscle fatigue, and prestrclching.
LENGTH-TENSION RELATIONSHIP
The force, or tension, that a muscle exerts varies
with the length at which it is held when stimulated.
This relationship can be observed in a single fiber
contracting isometrically and tetanically, as illustrated by the length-tension curve in Figure 6-9.
Maximal tension is produced when the muscle fiber
is approximately at its "slack," or resting, length. ,If
the fiber is held at shorter lengths. the tension falls
off slowly at first and then rapidly. If the fiber is
lengthened beyond the resting length, tension progressively decreases.
The changes in tension when the fiber is
stretched or shortened primarily are caused by
structural alterations in the sarcomere. Maximal
isometric tension can be exerted when the sarcomeres arc at their resting length (2.0-2.25 /J.m), because the actin and myosin fllamcnts overlap along
their entire length and the number of cross-bridges
is maximal. If the sareo meres arc lengthened. there
are fewer junctions between the marnents and the
active tension decreases. At a Sarcomere length of
approximately 3.6 J.Lm, there is no overlap and
hence no active tension. Sarcomcl·e shortening to
less than its resting length decreases the acLive ten-
Tension·length curve from part of an isolated muscle f
stimulated at different lengths. The isometric tetanic t
sion is dosely related to the number of cross-bridges o
the myosin filament overlapped by the actin filament.
tension is maximal at the slack length, or resting lengt
the sarcomere (2 p.m), where overlap is greatest, and f
to zero at the length where overlap no longer occurs
,....m). The tension also decreases when the sarcomere
length is reduced below the resting length. falling sha
at 1.6S IJ.m and reaching zero at 1.27 IJ.ffi as the exten
overlap interferes with cross-bridge formation. The str
tural relationship of the actin and myosin filaments at
ous stages of sarcomere shortening and lengthening is
portrayed below the curve. A. actin filaments; M. myo
filaments; Z, Z lines. Adapted from CriJ....,;ord. CN.C & J~
N. T. (1980). The design of muscfes. In R. Owen, J. Goodfelfo
P. Buffough (Eds.), Scientific Foundations of Orthopaedics an
Trmilllalology (pp. 67·-74). London.' William Heinemann;')5
modified from Gordon, A.M., rluxley, A.F.I., & JlIlian, F.l. (1
Tile variation in isometric tension with SMcomere length in
tebr~lre muscle fibers. J Physiol, 184, 170.
sian because it allows ovedapping of the thin
ments at opposite ends of the sarcomere, which
functionally polarized in opposile directions.
sarcomere length of less than 1.65 fJ.m. the thic
aments on the Z line and the tension dimi
sharply.
Thc length-tension relationship illustrated in
ure 6-9 is for an individual muscle fiber. If this
,lionship is measured in a whole mus.de contrac
isometrically and tetanically, the tension produ
by both active components and passive compon
must be taken into account (Fig. 6-10).
The cun/e labeled "active tension" in Figure
6-10 represents the tension developed by the contractile elements of the muscle. and it resembles
the curve for the individual fiber. The curve labeled
"passive tension" reflects the tension developed
when the muscle surpasses its resting length and
the noncontractile muscle belly is stretched. This
passive tension is mainly developed in the parallel
and series elastic components (Fig. 6-6). When the
belly contracts, the combined active and passive
tensions produce the total tension excrled. The
curve demonstrates that as a muscle is progl'cs·
sivdy stretched beyond its resting length, the pas·
sive tension rises and the active tension decreases.
Most rnuscles that cross only one joint normally
are not stretched enough fOl- the passive tension 10
play an important role, but the case is different for
lwo-joint muscles. in which the extremes of the
Eccenlric
Concentric
--~
u
'"o
-'
o
..
Velocity
load-velocity curve generated by plotting the veloci
motion of the muscle lever arm against the external
When the external load imposed on the muscle is ne
ble, the muscle contracts concentrically with maxima
speed. With increasing loads the muscle shortens mo
slowly. When the external toad equals the maximum
that the muscle can exert, the muscle fails to shorten
has zero velocity) and contracts isometrically. When
load is increased further, the muscle lengthens eccen
cally. This lengthening is more rapid with greater loa
L
Length
~------
I
I
The active and passive tension exerted by a whole muscle
contracting isometrically and tetanically is plotted against the
muscle's length. The active tension is produced by the con·
tractile muscle components and the passive tension by the series and parallel elastic components, which develop stress
when the muscle is stretched beyond its resting length. The
greater the amount of stretching, the larger the contribution
of the elastic component to the total tension. The shape of
the active curve is generally the same in different muscles.
but the passive curve, and hence the total curve, varies depending on how much connective tissue (elastic component)
the muscle contains. Adapted from Crawford. CN.C & James.
N. T. (1980). Tile design of muscles. III R. Owen. 1. Goodfel/o;,t~ & p.
Bulfough (£els.), Scientific Foundations of Orthopaedics and TraurnJ·
tology (pp. 67-74). London: William Heinemann.
I
•
length-tension relationship may be functi
(Crawford & James, 1980). For example, the
strings shorten so much when the knee is
Hexed that the tension they can exert decreases
siderably. Conversely, when the hip is nexed an
knee extended. the Illuscles arc so stretched t
is the magnitude of their passive tension tha
vents flirt her elongation and causes the knee t
if hip flexion is increased.
LOAD-VELOCITY RELATIONSHIP
The relationship between the velocity of shOl·
or eccentric lengthening of a muscle and dif
constant loads can be determined by plolLing t
locil)' of motion of the muscle lever arm a1 va
external loads. thereby generating a load-ve
curve (Fig. 6-11). The velocity of shortening
muscle contracting concentrically' is inverse
lated to the external load applied (Guyton,
The velocity' of shortening is grealest when th
~~~·_----_·_-------T
Gastrocnemius Muscie Tear
- -. ·.
A
22·year-old male professional athlete tears his ga5___ trocnemius during a race (Fig. (56-1-1). The tensile
overload that happens during strenuous eccentric and
c6n~entric contractions increases the risk of injury. especiallywhen the forces involve bj-articular muscles such
:asthe gastrocnemius. This indirect trauma is associated
:wi!~ 111gh tensile forces during rapid contraction (high
v~l_()~ity) and continued changes in muscle length. The
tcrnalload is zero, but as the load increases th
ele shortens morc and more slowly. \Vhcn the
nul load equals the mnximal force that the
can exen, the velocity of shortening becom
and the muscle contracts isometrically. \,\fl
load is increased still further, th~ muscle co
eccel1Lrically: it elongates during contractio
load~velocity relationship is reversed from
the concentri-cally contracting rnuscle; the
eccentrically lengthens more quickl)! with inc
load (Kroll. 1987) (Case Study 6-1).
status of muscle contracrion at the time of overload is
uS!Jally eccentric, and failure most often occurs at or
near the myotendinous junction unless the muscle has
Q.eef1. previously injured (Kasser, 1996). Swelling from
·h~~{br'rh~ge occurs
initially in the inflammatory phase.
,The cellu.lar response is more rapid and repair is (rlore
complete'if the vascular channels are not disrupted and
·thenutritioo.of ihe tissue is not disturbed. The degree
of injury from a tensile overload will dictate the poten·
tia/hos< r~sponse and the time needed iar repair.
FORCE-TIME RElATIONSHIP
The force, or tension, generated by a muscle
portional to the contraction time: the long
contraction lime, the greater is the force deve
lip to the point of maximurn tension. In Figur
this relationship is illustrated by a force-lime
for a whole muscle contl"acting isomet
Slower contraction leads to greaLcr force p
lion because time is allowed for the tensio
duced b.y thL' cornractilc clements to be trans
through the parallel clastic components to t
don. AltlH)Ugh tension production in the cont
component can reach a maximum in as little
mscc. lip to 300 ITlSeC ma~' be needed for th
sion to be transferred to the clastic compo
The tension in the tendon will reach the max
tension developed by the contractile element
the aClive contraction process is of sufficien
lion (Olloson, 1983).
EFFECT OF SKElETAL MUSCLE
ARCHITECTURE
The Illuscles consist of the contractile comp
the sarcomere, which produces active tensio
arrangement of the contractile components
the contractile propcnies of the muscle dr
cally. The more sarcomere lie in series, the
the myofibril will be; the more sarCOlllere lie
lel. the larger the cross-sectional area 01" the m
ril will be. These two basic architectural patte
myofibrils (long or thick) affect the contractile
erties of the llluscies in the following ways:
C~;~e-Study Fig~re 6-1-1,
;-" -.
1. The force the muscle can produce is pro
lion'll to the cross-section of the myofibr
(Fig_ 6-13A).
2. The velOcity and the excursion (working
range) that the muscle can produce are
portional to the length of the I11vofibril
(Fig. 6-138).
100
Muscles with shoner fibers and a Im'ger crosssectional area are designed to produce [orce,
whereas muscles \vith long fibers are designed for
excursion and velocity. The quadriceps muscle contains shorter myofibrils and appears to be specialized ror force product ion. The sartorius muscle has
longer fibers and a smaller cross-sectional arca and
is better suited for high excursion (Barana et aI.,
1998; Lieber & Bodine-Fowler, 1993),
80
~
~
~
60
~
40
~
::;
"
20
EFFEa OF PRESTRETCHING
10
It has been demonstrated in amphibians and in humans (Cuillo & Zarins, 1983) that a muscle performs more work when it shortens inlmediatcly af-
15
20
Muscle Length (cm)
100
B
ter being stl'etched in the concentrically contracted
state than when it shortens from a state of isometric contraction. This phenomenon is not entirely accounted for by the elastic energ~' stored in the series
clastic component during stretching but must also
be caused by energy stored in the contractile com·
80
~
~
o
60
u.
~
'0
~
40
::;
"
20
5
10
15
20
Muscle Velocity (cm/s)
Isometric and isotonic properties of muscles with differ
architecture. A. Force-length relationship. B. Force-veloc
relationship. PSCA. physiological cross-sectional area.
Reprinted with permission of the AmeriC<Jn Physical Therapy
sociation from Lieber, RL (J 993). Skeletal muscle rnec!wnics
Implications for rehabilitation. Physical Therapy, 73(2). 852.
•
0F-----------
pon~nl. 'It has been suggested [hat changes in the
trinsic mechanical properties of l11)'oflbrils are
ponant in the stretch-induced enhancement
work production (Takarada et aI., 1997),
Time
IBm
i
L-
_
Force-time curve for a whole muscle contracting isometri·
cally. The force exerted by the muscle is greater when the
contraction time is longer because time is required for the
tension created by the contractile components to be transferred to the parallel elastic component and then to series
elastic component as the musculotendinous unit is
stretched.
EFFECT OF TEMPERATURE
A rise in muscle temperature causes an incre
in conduction velocity across the sarcolem
(Phillips & PeLrorsky, 1983), increasing the
quency of stimulntion and hence tbe product
of muscle force. Rising of the muscle temperat
from 6 La 34 u C results in an almost linear incre
or the tension/stirrness raLio (Callcr et al" 199
A rise in temperature also causes greater enzymatic activity of muscle metabolism, thus increasing the efficiency of rnusclc contraction. A
further effect of a rise in temperature is the increased elasticity of the collagen in the series and
parallel elastic components. which enhances the
extensibility of the muscle-tendon unit. This increased prestrclch incrcases the force production
of the muscle.
Muscle temperature increases by rneans of two
mechanisms:
Fatigue in a muscle contracting isometrically. Prolong
stimulation occurs at a frequency that outstrips the m
cle's ability to produce suHicient ATP for contraction.
result, tension production declines and eventually ce
Adapted from Luciano. 0.5., Vander. A.i., &- Sherman. J.H
(1978). Human Function and Structure (pp. 113-136). Ne
I. Increase in blood now, which occurs when an
athlete "warms up" his or her muscles
2. Production of the heal of reaction generated
by metnbolism. by the release of the energy
of contraction, and by friction as the contractile components slide over each other
However, at low tenlperaturc (1 O('C), it has been
shown that the maximum shortening velocity and
the isometric tension are inhibited significantly.
This is caused by decreased pH (acidosis) in Ihe
muscle. The 1'1-1 plays a much less important role at
temperatures close to the physiological level (Pate et
aI., 1995).
EFFECT OF FATIGUE
The ability of a muscle to contract and relax is dependent on the availability of adenosine triphosphate (ATP) (Box 6-1). If a muscle has an adequate
supply of oxygen and nutrients that can be broken
down to provide ATP, it can sustain a series of lowfrequency twitch responses for a long time. The rrequency must be low enough to allow the muscle to
synthesize AT? at a rate sufficient to keep up with
the rate of ATP breakdown during contraction. If
the fTequency of stimulation increases and outstrips
the rate of replacement of ATP, the twitch responses
soon grow progressively \veaker and eventually fall
to zero (Fig. 6-14). This drop in tension following
prolonged stimulation is muscle fatigue. If the frequency is high enough to produce tetanic contractions, fatigue occurs even sooner. [f a period of rest
is allowed before stimulation is continued, the ATP
concenln.l.lion rises and the muscle briefly recovers
its contractile ability before again undergoing fatigue.
Three sources supply ATP in muscle: creatine
phosphate, oxidative phosphorylation in Ihe mitochondria. and substrate phosphor:ylation during
anaerobic glycolysis. \'Vhen contraction begins, the
York: McGraw-Hili.
•
myosin ATPase rapidly breaks down ATP. Th
crease in adenosine diphosphate (ADP) and
phate (Pi) concentralions resulting from this b
down ultimately leads to increased rale
oxidative phosphorylation and glycolysis. A
short lapsc. however. lhese metabolic palhway
gin lO deliver ATP at a high rate. During this
vaL lhc energy for AT? formation is provided b
atine phosphate, which offers the most rapid m
of forming ATP in the muscle cell.
At moderate rates of muscle aClivity, most o
required ATP can be formed by the process o
idalive phosphorylation. During intense exe
when ATP is being broken down rapidly, the
abilitv to replace ATP by oxidalive phosphory
may be limited. primarily by inadequate deliv
oxygen to the muscle b.y lhe circulatory system
Even \vhcn oxygen delivery is adequate, th
al which oxidative phosphorylation can pro
ATP may be insufficient to sustain inlense ex
because the enzymatic machinery of this pathwC
relatively slow. Anncrobic glycolysis then beg
contribute an increasing portion of the ATP. Th
colytic pathway, although it produces much sr
amounts of ATP h-om the breakdown of glucos
erates at a Illuch faster rate. It can also proce
the absence of oxygen, with the formation of
acid as ils end product. Thus, during intense
cise, anaerobic glycolysis becomes an addi
source for rapidly supplying the muscle with A
The glycolytic pathway has the disadvanta
requiring large amounts of glucose for the pr
tion of small amounts of" ATP. Thus, even th
muscle stores glucose in the form of glycogen,
ing glycogen supplies rnay be depicted quickly
muscle aconty is intense, Finally, myosin ATPase
may break down AT? faster than even glycolysis can
r~p'lace it, and fatigue occurs rapidly as AT? concentrations drop.
After a period of intense exercise, creatine phosphate levels have become low and much of the muscle glycogen may have been convened to lactic acid.
For the muscle to be returned to its original state,
creatine phosphate must be resynthesized and the
u1vcofLcn stores must be replaced. Because both
pl:oce~ses require energy, the muscle will continue
to consumc oxygen at a rapid rate cven though it
has stopped contracting. This sustained high oxy!:rell uptake is demonstrated by the fact that a person
~ontinues to breathe heavily and rapidly aher a period of strenuous exercise.
v"hen the energy necessary to return glycogen
and creatine phosphate to their original levels is
taken into account, the ef(-icienc~. . with which muscle coovens chemical energy to work (movement) is
usually no more than 20 to 250/0, the majority of the
energy being dissipated as heat. Even when Illuscle
is operating in its most efficient stnte, a maximum
()f only approxinwtel)! 45{>~ of the energy is used for
contraction (Arvidson ct 'II., 1984; Guyton, !986).
In growth biomechanics, muscle fatigue is flrst
observed by the lack of coorclinalion or movement
and its effect in the increasing of loads in tissue. Researchers including Bates et al. (1977) have indicated that the skill of the person in performing a
given action is affected by fatigue, They studied the
fatigue effect on runners and absented that runners
decrease their knee extension when fatigue occurs
(Bates et aI., 1977). Parnianpour (1988) studied the
motion coupling of the spine at exhaustive extension
Oexion. This study showed that when an individual
became fatigued, the coupled motion increased and
therefore the spinaltorquc increased. The most deleterious component of the ncurornuscular adaptation
to the fatiguc state was the reduction in accuracy
control and speed of contraction, \vhich may predispose an individual to injury if muscle fatigue occurs.
Muscle Fiber Differentiation.
In the preceding section, we described the major
ractors that determine the total tension developed
b.v the whole rnuscle when it COlllracts. Individual
muscle fibers also display distinct differences in
their rates of contraction, development of tension.
and susceptibility to fatigue.
Jvlany methods of classifying muscle fibers
been devised. As early' as 1678, Lorenzini obse
anatomically the gross dilTcrence bel\\,Icen red
white muscle. and in 1873 Ranvier typed muscl
the basis of speed of contractility and fatigab
Although considerable confusion has existed
cCl'ning the method and terminology for classif
skeletal muscle, recenl histological and histoch
cal observations have led to the identificatio
three distinct typcs or muscle fibers on the bas
differing contractile and metabolic prope
(Brandstater & Lambert, 1969; Buchtah! & Soh
burch, 1980) (Table 6-1).
The fiber t~'pes are distinguished mainly by
metabolic pathways by which they can gene
AT? and the rate at which its energy is made a
able to the conrractilc system of the sarcom
which determines the speed of contraction.
three fiber types are termed type I, slow-twitch
idative (SO) fibers; type lIA, fast-twilCh oxida
glycolvtic (FOG) fibers; and type liB, fast-twitch
colytic (FG) fibers.
Typc I (SO) fibers are characterized by a low
tivity of my'osin ATPase in the muscle fiber
thcrefore, a relatively slow contraction time.
glycolytic (anaerobic) aClivity is low in this
lype, but a high content of mitochondria produc
high potential for oxidative (aerobic) activity. Ty
fibers are difficult to fatigue because the high ra
blood now to these fibers delivers oxygen and n
ents at a sufficient rate to keep up with the relati
slow rate of ATP breakdown by myosin ATP
Thus, the fibers are well suited for prolonged,
intensity work. These fibers are relatively sma
diameter and so produce relatively little tens
The high myoglobin content of type I fibers give
muscle a red color,
Type II muscle fibers arc divided into two n
subgroups, IIA and IlB, on the basis of differing
ceptibility to treatment with different buffers p
to incubation (Brooke & Kaiser, 1970). A third
group, the typc lIe fibers, are rare, undifferenti
fibers, which are usually seen before the 30th w
of gestalion. This fiber type is infrequent in hu
muscle (Banker, 1994). Tvpe IIA and 11 B fibers
characterized by a high activity of myosin ATP
which results in relatively fast contraction.
Type lIA (FOG) fibers are considered interm
ate between type 1 and type lIB because their
contt-action time is combined with a modera
well-developed capacity for both aerobic (ox
tive) and anaerobic (glycolytic) activity. T
Properties of Three Types of Skeletal Muscle Fibers
TYPE 1
Slow-Twitch
Oxidative
TYPE JlA
Fast-Twitch
OxidativeGlycolytic
(SO)
(fOG)
Glycolytic
(fG)
Speed of contraction
Slow
Fast
Fast
Primary source of ATP
production
Oxidative
Oxidative
phosphorylation
Anaerobic
phosplloryla tion
glycolysis
Glycolytic enzyme activity
low
Intermediate
High
Capillaries
Many
Many
few
Myoglobin conlent
High
High
low
Glycogen content
Low
Intermediate
High
Fiber diameler
Small
Intermediate
Large
Rate of fatigue
Slow
Intermediate
Fast
fiber's also have a well-developed blood supply.
They can maintain their contractile activity for rel<Hively long periods; ho\vevcr, at high rates of activity. the high rate of ATP splitting exceeds the capacity of both oxidative phosphorylation and
glycolysis to supply ATP, and these fibers thus
eventually fatigue. Because the myoglobin conlCIH
01" this muscle type is high, the muscle is of len cat-
egorized as red III lIscl e.
T:vpe liB (FG) libers rely primarily on glycolvtic
(anaerobic) activity for ATP production. Few capillaries are found in the vicinity of these fibers and because they contain little myoglobin (hey are ohen
referred to as white muscle. Although Lype liB fibers
are able to producc ATP rapidly, they fatigue easily
because their high rate of ATP splitting quickly depletes the glycogen needed for glycolysis. These
fibers generally are of large diameter and are thus
able to produce great tension, but only for short periods before they fatigue.
It has been well demonstrated that the nerve innervating the muscle fibcr determines its t.ypc
(Burke el aI., 1971); thus, the muscle fibers of each
motor unit are of a single type. In humans and other
species, electrical stimulation was found to change
the fiber type (Munsal, McNeal, & Waters, 1976). In
animal studies, transecting the nerves that innervate slow twitch and fast-twitch muscle fibers and
then crossing these nen;es was noted to reverse the
fiber types. After recovery from the cross-innervation, the slow-twitch fibers became fast in their con8
TYPE JIB
Fast-Twitch
tractile and histochemical properties and the
twitch fibers became slow.
The fiber composition of a given muscle
pends on the function 01" that muscle. Some m
cles perform predominantly one form of cont
tile activity' and arc often composed mostl
one muscle fiber t}'pc. An example is the so
muscle in the calf, which prirnarily maint
posture and is composed of a high percentag
type I fibers. More commonly', however, a mu
is reqllir'cd to perform endurance-type act
unde)' some cirClllTlstances and high-inten
strength activity under others. These mus
generally contain a mixture of the three mu
fiber types.
(n a typical mixed muscle exerting low tens
some of the small motor units, composed of ty
fibers, contract. As the muscle force increa
more motor units are recruited and their
quency of stimulation increases. As the frequ
becomes maximal. greater muscle force
achieved by recruitment of larger motor u
composed of type 11;\ (FOG) fibers ancl eventu
type lIB (FG) fibers. As the peak muscle force
creases, the larger units are the first to cease
tivity (Guyton. 1986; Luciano, Vander. & Sherm
1978).
It is generally, but not universally, accepted
fiber types are genetically det~rmined (Costi
aI., 1976; Gollnick, 1982). In the average pop
tion, approximatel.'Y' 50 to 55% of muscle fi
are type I, approximately 30 to 35 Cj(; are type IrA,
and approximately' 15 rJr; are t.\'Pe lIB, but these
percentages vary greatly among individuals.
In elite athletes, the relative percentage of fiber
types differs from that in the general population and
appears to depend on whether the athlete's principal
activit)' requires a short, explosive, maximal effort
or involves submaximal endurance. Sprinters and
shot putters, for example, have a high percentage of
tvpe II flbers, whereas distance runners and cross~ountry skiers have a higher percentage of type I
fibers. Endurance athletes may have as man.v as
80% t.ype I fibers, and those engaged in short, explosive efforts as few as 30 0/e of these fibers (Saltin
et 'II., 1977).
The genetically determined fiber typing may be
responsible for the natural selective process by
which athletes are drawn to the type of sport for
which they are most suited. Because fiber types are
determined by the nerve that innervates the muscle
fiber, there may be some cortical control of this innervation that influences an athlete to choose the
sport in which he or she is genetical I)' able to excel.
Muscle Injuries
iVluscle injuries comprise contusion, laceration,
ruptures, ischemia, compartment syndromes, and
denervation. These injuries weaken the muscles and
can cause significant disability'. Blunt trauma can
diminish muscle strength, limit joint motion, and flnally lead to myositis ossiflcans. lVlusclc laceration,
surgical incisions, and traumatic lesion to muscle
tissue and denervation weaken the ITmsclcs, sometimes significantly. Ruptures in muscles also can
cause weakness. Like the other injuries, they rnay
result from direct trauma, but muscle contractions
against resistance also can lead to tears in muscle
tissue.
Acute muscle ischemia and compartment syndromes can cause extensive muscle necrosis. The
many potential causes of compartment s}'ndrome
all result in increased pressure \vithin a confined
muscle compartment. In this case, failure to relieve
the pressure rapidl.y ma}' cause complications that
range from weakness and decreased motion to loss
of an entire limb.
Studies have shown that healthy skeletal muscle
has a substantial capacity to repair itself. This repair process following a specific injury is inferred
by the prior innovation pattern, vascularization,
physical constraint of the surrounding tissues, t
extent and condition of extracellular matrices, a
the development of repair cells. Muscle injuries a
important but the topic is not \vithin the scope
this chapter. Injuries should be investigated ca
fully if suspicion arises that a patient has mus
damage.
Muscle Remodeling
The remodeling of muscle tissue is similar to that
other skeletal tissues such as bone, articular car
lage, and ligaments. Asin these other tissues, mu
cle atrophies in response to disuse and immobili
tion and hypertrophies when subjected to grea
use than usual.
EFFECTS OF DISUSE AND
IMMOBILIZATION
Disuse and immobilization have detrimental effe
on muscle ftbers. These effects include loss of
durance and strength and muscle atrophies on a m
crostructural and macrostructural level, such as
creased numbers and size of fiber. Biochemi
changes occur and affect aerobic and anaerobic
erg}' production. These effects are dependent
fiber type and muscle length during immobilizati
Immobilization in a lengthened position has a l
deletet-ious effect (Appell, 1997; Kassel', 1996; Oh
et 'II.. 1997; Sandmann, et 'II., 1998).
Clinical and laboratory studies of human and
imal muscle tissue suggest that a program of
mediate or early motion may prevent Illuscle at
phy after injury or surgery. In a study of cn
injuries to rat muscle, the effect of immobilizat
of the crushed limb was compared with that of
mediate motion. The muscle fibers \vere found
regenerate in a more parallel orientation in the m
bilized animal than in the immobilized anim
capillarization occurred more rapidly, and tens
strength returned more quickly. Similar resu
were found in a later study on the effect of imm
bilization on the IT'lOrphology of rat calf musc
(Kannus et aI., 1998a).
It has been found clinically that atrophy of
quadriceps muscle that develops while the limb
immobilized in a rigid plaster cast cannot be
v'ersed through the use of isometric exercises.
rophy may be limited by allowing early mot
such as that permitted by a partly mobile c
brace. In this case, dynamic exercises can be performed.
Human muscle biopsy studies have shown that
it is mainly' the type I fibers that atrophy with immobilization; their cross-sectional area decreases
and their potential for oxidative enzyme activity is
reduced (Kannus et aI., 1998b). Earl)' motion may'
prevent this atrophy. It appears that if the muscle
is placed under tension when the body segment
moves, afferent (sensory) impulses from the intrafusal muscle spindles will increase, leading to increased stimulation of the type I fiber. Although intermittent isometric exercise may be sufficient to
maintain the metabolic capacity of the t.vpe II
fibel~ the type I fiber (the postural fiber) requires a
more continuous impulse. Evidence also suggests
that electric stimulation may prevent the decrease in type I fiber size and the decline in its oxidative enzyme activity caused b:v immobilization
(Eriksson et ai., 1981).
In elite athletes, inactivity following injury,
surgery, or immobilization rapidly decreases the
size and aerobic capability of muscle fibers, particularly in the fiber type affected bv the chosen
sport. Tn endurance athletes, type I fibers are affected, while in athletes engaged in an explosive
activity such as sprinting, type II fibers are affected.
stood (Gollnick, 1982; Guyton, 1986). It ap
that these evcnts arc controlled or modifi
both the intrafusal muscle spindles, located
allel with the extrafusal fibers of the muscle
and the Go!gi tendon organs, located in serie
Ruptured Left Anterior eruciate Ligam
A
25-year-old male, status postsurgical repair of the
- ruptured left anterior cruciale ligament, had torqu
measurements taken from the involved and uninvolved
limb 10 weeks after the surgical procedure (Fig. (56-2-1
and repeated 6 weeks after the training began (Fig. (56
2-1 B). An increase of muscle torque is shown in the repeated isokinetic test. The initial deficit of the involved
side was approximately 63~o when compared with the
uninvolved side. After 6 '.Neeks of trainin~J. the deficit o
the involved side compared with tho uninvolved side de
creased to 43%.
187,------------163
140
E 116
<n
c
93
;:o 47
"
z
EFFECTS OF PHYSICAL TRAINING
Physical training increases the cross-sectional area
of all muscle fibers, accounting for the increase in
muscle bulk and strength. Some evidence suggests
that the relative percentage of fiber types composing a person's muscles rna:v also change with physical training (Arvidson, Eriksson, & Pitman, 1984).
The cross-sectional area of the fibers affected by
the athlete's principal activit).' increases. For example, in endurance athletes, the arca of muscle taken
up by type I and type IIA fibers increases at the expense of the total area of type lIB fibers (Case
Study 6-2).
Stretching incrcases muscle Oexibility, maintains
and augmcnts the range of joint motion, and increases the elasticity and length of the musculotendinous unit (Brobeck, 1979; Cuillo & Zarins,
1983). It also permits the musculotendinous unit to
store more energy in its viscoelastic and contractile
components.
The events that take place during muscle
stretching are complex and incompletely under-
/
69
23
-----
o J-~c_-----__:-~:-''''-'---o 16 32 48 64 80 96 112
Knee motion degrees
A
294
239
204
~c 192
0
;: 136
108
z 68
33
0
" "
/.
.<
"
"-
"-
/
0
16
32
48
64
80
96
,
112
Knee motion degrees
B
Case Study Figure 6-2-1 Isokinetic test at
180~/sec.
A
Measurement of the quadriceps femoris torque produc
tion at 10 weeks postsurgical procedure. The dashed fi
represents torque output by the involved limb. The so
line represents torque output by the noninvolved limb
B. Measurements of the quadriceps femoris torque pro
duction at 16 weeks postsurgical procedure and 6 wee
after training sessions. The dashed line represents torq
output by the involved limb. The solid line represents
torque output by the noninvolved limb.
these fibers. The spindles respond to an increase in
muscle length and the Golgi apparatus to an increase in TllUSc!C tension. The resulting spindle reflex increases muscle contraction, while the Goigi
reflex inhibits contraction and enhances muscle
relaxation.
The intrafusal muscle spindles are or t\Vo types:
primary and secondary. The primary spindles respond to changes in the rate of muscle lengthening (dynamic response) and the actual amount or
lengthening. The secondary spindles respond only
(0 the actual length change (static response). The
static response is weak and the dynarnic response
is strong: therefore, keeping the rate of stretch
low may allow the dynamic response to be bypassed, essentially negating dte effect of the spindles. Conversely, the increase in rnusclc tcnsion
during stretching may activate the relaxing effect
of the Golgi apparatus and thus enhancc further
stretching. The various 111ethods and theories of"
stretching all have as a common goal inhibition of
the spindle effect and cnhanccrnent or the Golgi
effect to relax the muscle and promote further
lcngt hen i ng.
Sllmm.ary
i!;< The structural unil of skeletal muscle is the
fil.;·c'r, which is encompassed by the endomysium
and organized into fascicles encased in the perirnyslum. The epimysium surrounds the entire musclc.
2 The fibers are composed of myofibrils. aligned
so as to create a band pallern. Each repeat of this
pattern is a sarcomere, the functional unit 01" the
contractile system.
3 The myofibrils arc composed or thin filaments
of the protein actin and thick filaments of the protein myosin, and the intramyofibrillar cytoskeleton
is composed 01" the clastic filaments titin and the inelastic filarnents ncbulin.
4· According to the sliding filament theOl)" active
shortening of the muscle results from the reiativL'
movement of the actin and myosin filaments past
one anothel: The force of contnlClion is developed
by movement of the myosin heads, or cross-bridges,
in contact with the aclin filaments. Troponin and
tropomyosin, two proteins in the actin helix, reguIatc the making and breaking of the contacts between r-ilamenls.
5 A key (0 (he sliding mechanism is the calciu
ion, which lu)-ns the contractile activity on and a
6 The mOlOr unit, a single motor neuron and
musclc fibers inncrvatcd by it, is the smallesl pan
the muscle thal can contral:t independently. T
calling in of additional motor units in response
greater stimulation of thc rnotor nerve is known
recruitment.
7 The tendons and the endomysium, perim
sium, sarcolemma, and epimysium reprcsent par
lel and series clastic components dUll stretch w
aClive contraction or passive muscle extension a
recoil with muscle relaxation.
8 Summation occurs whcn mcchanical
sponscs of the muscle to successivc stimuli a
added to an initial response. \Vhen maximal tensi
is sustained as a result of summation, the mus
contracts tetanically. The muscle fiber contracts
an all-or-nothing fashion.
9 Muscles may contract concentrically, eccent
cally, or isometricaH)' depending on the relationsh
belween the muscle lension and thc rcsistance to
overcome. Concentric and eccentric contractio
involve dynamic work, in which t he muscle move
joint or controls its 111ovcn1cnl.
',10 Force production in muscle is influenced
the length-tension, load-velocity. nne! rorce-time
lationships of the muscle. The length-tension re
tionship in a whole Illuscle is influenced by both
tive (contractile) and passive (series and paral
elastic) components.
11 Two other factors that increase force produ
tion are prestretching or the muscle and a rise
muscle tempenllure.
12 The energy for muscle contraction and its
lease is provided by the hydrolytic splitting of AT
Muscle fatigue occurs \\'!hen the ability of the mu
cle to synthesize ATP is insufficient to keep
with the rate of ATP breakdown during contra
tion.
13 Three main fiber types have been identifie
type I, slow-twitch oxidative; type IIA, fast-twit
oxidative-glycolytic; and type liB. fast-twitch g
colytic fibers. Most muscles contain a mixture
these types.
.,14 Muscle atrophies occur under disuse and i
mobilization; muscle trophism can' be restor
through early and active rcmobilization.
REFERENCES
Appell, H.J. (1997). Thl' muscle in the rch'lbilit~l!ion pro("e~s.
OnJlOp{/{!t'. 26( 11).930-934.
AITidson. I., Eriksson. E.. 6.:. Pitman, M. (1984). Neuromuscular basis of n:.. habilit'llioll. In E. Hunter & J. Funk (Eds.),
RelllI!JiUtarioll
the "ljlt/wl I\."m'c (pp. 210-234). SI. Louis:
or
C. V. Mosby.
B;:lnkcr, B.O. (1994). Basic reaction of 1ll1lSt:!t:. In A.G. ErH!cl
& C. Franzini-Al'fllstrong (Eds.), Myology (2nd cel.). N~\\'
)'ork: McGraw-HilI. Inc.
Baratta R.V., Solornono\\', \1., Z!lOtl, S.H. (1998), Frequency
domain-based llwdds of skeletal llluscle. J ElcclroJllyogr
Kilu:siol, 8(1}, 79·91.
Billes. B.T.. Ostl,.'l'lling. L.R., James. S,L. (1977). Fati!!uc ef·
feels in running. ) MOlor Be/ull'. 9. 203-107.
-
Brobeck, J. R. (Ed.) (1979). Bes/ lIml Ttl.do,.:, Physiological Bt/·
sis of :\let/iclIl Practice (10th cd., pp. 59-113). B~llimore:
\\'illinms &. Wilkins.
Brooke. ;\'1.1-1. &. K~iscr, K.K. (1970). Thn:e lllyosin adenosine
Iriphosphat~sl: systcms: The nature of t1~cir pH liability
and sulfhydryl depi.'nd"llce. J l-listodl/!JIl Cy/oclli:lll. 18,
670.
Blll;ht~lhJ.
F & Sohl1li.\lburch, H. (1980). Motor unilS of Ill;llll11l;llian Illusdt.'. Phniol Ret'. 60. 90.
Burh-, R.E .. Levinc, D.N.. Z~lj~~C. F.E. (1971). M~lnlllwlinn mo·
tor unils: Physiological hislochemic~ll corrckltion in Ihrce
types of motor units in cal gastrocllemius. Seicllct.', 17-1,
709.
Cost ill, D.L.. Coyle, E.F., Fink. W.F., l.. eslllcs, G.R .. Wilzmann.
EA. (1979). Adapt.uioll in skeletal llluscles following
strengLh Lraining. J .-Ippl Pln'siol, 46.96·9,
~
Craig. R. (199-4), The structur~ of (he contract filaments. In
A.~. ~ngd ,&. C. Fr<lnzini-Annstrong (Eds.), ,\(mlogy (2nd
t.'cI.). New )ork: i\1cGraw·Hill, Inc
Crawford. C.N.C. &. J<101es. N.T. (1980). The dl'sign of muscles .. In R. O""..'n, J. Goodfellow, ~ fl. Bullollgh (Eds.) Sci·
~1!.lilie FOtllldaliOll.'
Onltopacdics (lIId Tnl1lll/tllO!ogy (pp.
0/-74). London: Willi.lm Heinemann.
Cuillo, J.V. & Zarins, B. (1983). Bioml,.'chanics of the mLlSt:Ulotcndinous unit: Relation 10 athk-lic performance and in·
jury. Clilt Sports ,\lcd, 2. 71.
Eriksson, E., Hagglll;uk, T, Kiessling, K.H., l,.'t;:\1. (1981). EffeCI of dectrie:ll slimulation on human skelt.-Ial musde.
1m Sports :lIed. 2, IS.
Galler, S.. Hilbcr. K. (1998). Tension/stirrncl-is ratio of skinned
rat skelcLallllllsde fibn.' types at v:\l'iolls lCl11pCratllrt's. Ac/
Physio! Sctlml, 162(2), 119-126.
Gollnick, P.O. (1982). Ri..'lationship or strcngth and cndllr<lIlce
with skeletal muscle structure and llll,.·tabolic potential. lilt
J Spons ,\1t:tI (SliP!'!]. 3, 26.
Gordon, A.M., Huxley, A.F.!., &. Julian, FJ. (1966). The variation in isol1\l,.'tric 1i..'nsiOll with s;trC<llllt.'rt.' kngth in \'I:.'rtebratc Illuscle fibers. J Physiol, 18-1, 170.
Guylon, A.C. (1986). Tex/book
,\ledieal Pltysiolog\' (7th e(l.).
Philndelphia: \V.B. Saunders.
.
Hnm,. A.W. &. Corm:lck, D.H. (I979). Nis{olo~y (Sth cd.).
Phl1adclphia: J.B. Lippincott.
Hill, A. V. (1970). Firsl (//Ill L{lsi £'\"/1(:ri1l11:11/:; ;11 ,\fl/sell! ,lh'·
dUlllies. Cambridgl..': C;imbridgl: Uni\'cl"sitv Press.
Hill, O.K. (1968), Tl.:nsion due 10 interaction bt.~lween lhe sliding filal1lt.'nts of resting $lriatcd museit.-. Thl,.' dft.'cr of
stimulatinn. Pltysio! (LOllll], /99,637.
or
or
-'
Hi!ilop, H.J. ~ Perrine, J. (1967). The isokinelic I.:Olltl.'p
cn:isc. Ph."..; Ther, -Ii. 114.
Huxley, A.F (1974). Muscul"r contraction. J Physi<:d, 2Huxley. A.F &. Huxll.'Y, H.E. (196-1). Organizcrs of:t disc
of Ihe physical and chemical basis of l1l11st.:lllar tO
tion. Pmc
Sec. 8160. -433.
Kannus. P., JOI.5", L.. K\·j:.;1. ,\1., Jarvinen, T" J;lninc
(1998aL EHeus of immobilizaLion and sllbseqm'n
and high. intensity excrcise on morphology of I'~\t tal
cles. Scal/t! J ,\Jed Sci Sports. 8(3), 160-171.
Kant1us. P., Jozsa, L.. Jar\'int'n, T.L., K\'ist. 1\'1.. Vie
Jilrvinen, ~\'t. (1998b). Fn:e mobilization ,11ld low- to
intensity exercise ill immobili/.ation-indllccd muscl
phy. J ..\ppl P/I\'s;ol, 8-1(4), 1418-1424.
Kassel', l.R. (1996). Celleral Knowledge. III J.R. Ka5sI..
Onltop{lcdic Kllou·l.... d~i! lIpt/tltf! 5: I-Ioml.' Sr/flh- S\'
Illinois: American :\cad('my of Orthopaedic Su;·!!c(;
Kedc. C.r\ .• Nl.'il. E., ..\: Jock N, (1982). \Iusch: a,;d Ih
vous system. In StllJl:WI/ W,.ighl~ Applied Phrsiolo"\"
cd .. pp. 248-259). Oxford: O.xford Unin.'l'siL\: Pres~~'
Komi, P. V. (1986). Tbl,.' slrl,.'tch-shorti.'nin~ t:v~le and
power OUlpUt. In N.L. JOllCS, N. McCan·!lt.'\·, ~ A.
Con~s (Eds.) I-ItlIIltlJl "'//SciI.' POII'a (PI'. 2'7-39).
paign. IL: 1-lull1an Kin .... tics Publishers.
r\.I'Oellll,.'l', I\..H.E., Marr:ls. \V.M., McGlothlin, J.D.,
(1990). On the IlH::ISUl'l..'lllent of hllm~\ll stf'I..'Il!.nh. III
elliS! [;'/",!!,OIlOlllics, 6, 199-21 O.
~
Kroll. P.G. (1987). The <fjfi:Cl oFpl"t.'l'iolls co!//rac/ioll cO
011 SU!lS(,({IIl'11I l..'cn'Jllric PO\\'('I" production iu dhol\'
IIIIISC!CS. Unpllbli~lwd doctoral dissertaLioll, New
University. Nt::\\· York.
Lkbcr, RL. &. Bodine·Fowler, S.c. (1993). Sh'Il,.,t~d 1ll1lS
~hanics. ImpliGllions for n.'hahilit<ltion.l'h.l's Tho:
n
844-856.
Linke, W.A .• In:·l1lcyl,.'l', ~L '\-llInde!. P., Stockm,·icl'.
Kolml,.·rL'r, B. (1998). N~lturl,.· of PEVK-titill e1asli
skclet~ll muscle. PJ'(l(: NIlt! '-\1.:lId Sci liS..\.
8052-8057.
Luciano, D.S., Vander, :\.1 .. .& Sherman, J.H. (1978). I
I-'Il.I/clio/1 alll! Strl/('ltlrt' (pp. 113-136). New York: \k
Hill.
\'lul1s.al. TL. NcNl.'al. D.. & Wat(-rs. R. (1976). Eff
lll,.'l"\,e stimulation on human must-Ie. :lrcll N<'lI.rol, 3
Ohir~l. Y.. Y~\slli, w., Roy, R.R., Edgerlon, \f.R. (199;).
of mllscl\: ll:ngLh on 1111: r('spons\: to unloadirH2. ..le!
([Jasen. 159(2-3), 90-98.
Ottoson, D. (1983). I'hysio!o.!,:y of lhi' NO"l'IllIS 8"SI<,1
78-116), New York: Oxford Univcrsitv Press.
.
Parnianpou[", ~...I.. Nordin, M .. Kahano\"iLz. N.. "'I al.
The triaxial collpling torque generation of trunk m
during isometric I..'xl,.'nioI1S and tlti.' l,.'fb.:1 of fat
isoinerlial mOH'ltIents on the molor output <lnd mOH
pallcrns. Spillt.', 13(9).
Pate, E., Bhimani, i\l., Fnmks-Skib'l, K., Cook. R. (199
duced ('ffCCI of pH on skinned r~lhbit jlSO:lS musc
chanics at high temperatures: Implications for f"l
P!lysiol (Loud), -I86(Pt 3), 689-694.
Phillips, C.,x. &. Pctrofsky, 1.5. (1983) . .\lrchfUlics of S
alltl e(mliut.' ,\fuse/e. Springfield; Charles C. Thon1<ls
S"ltin. B.• el al. (1977). Fiber types and I1K'tabolk polt
of skdelal Illuscles in sedcntal"}' m;m and enduranc
n(;rs. :11111 NV ..lead Sd, 301. 3.
Snndm<tnn, ~I.E .. Shoemanll. J.A .. Thompson. L. V, (1998).
Th..:- fib1.'r·typ ...·spccific effcct of inactivity and intcrmillellt
wci2ht·henring on the !!aslroc!lcmius of 30'lllollth-old
rats~ Arch PI/\'~ ,lied UCllllb;/, 79(6), 658-662.
Squire. J.~1. (1997). Architecture and fUllction in thl.: muscl~
s~lI·comcrl·. e/ln Opill Stl'llCI Bio!. 7(2). 247-257.
Strom~r, M.H. (1998). Thc cytoskelelon in skckt;-.l. c:lI'di:1C
:lIId smooth muscl~ (dis. His/o! f/islOpa/llOl, / 3( I).
283-291.
-'::' '
'-'"
i.:
",'
'''"
T:lbrad~l,
Y., (wnll\oto, 1-1., Sugi. 1-1 .. Hir~ln(), Y.. Ishi
(1997). Strctch-ilH.luccd l'nhanCCIllc.'Jll of m':"chanical
production in long frog single fihers and hUIll<ln mus
App! Physio!, 83(5),1741-17-4$.
Wilkie. D.R. (J956). The mechanical propenies of llluscl
.\It'd BIIII, /2, 177.
Wilkie, D.R. (1968) . .\lusc/c. london: Edward Arnold.
Williams, P. &. Warwick. R. (1980). Gray:.; rlJlll/om." 06tl1
pp. 506-5(5). Edinburgh: Churchill Livingstone.
z
0
;=
--:-----
~
---------1
e-
o
Z
.~
0
u
~
1<
>!'
z
"
0
0
"9
:>
~
0
>
~
v
'"
E
<;;
, ....
0;
0;
,.'
.,;
~
'"
(;
c
.2
:;;
.~
c
,·_------·1·----------1'
',." ,.
.
~
.
'"
<;
-0
.
··
·-~
~
-~=
6
'g"
.~
.~
"=c
:,
<;
=
<;
e
t
~
~
2
~
~
<
§
"=
0
.~
"
.~
",
-5
g
;::
........ -
..
"
<;
'"
l'
C
,
"=
~I
.
~
c.
,~:':,i$;\')
E
.2X
;;;
v
.a
;1~&
0
c
~
:s
.:)~;;;:
'"'"
E
,.', u....
'.;'x.
~
j,t-:;:".
. e2?
~
<~-
ij
X
~
",:tt:./
s"
.~
."
~
v
~
~
E
-S
.~
0
0
£
0
0
E
-0c
.-
"
.;;
0
3
0
;;:
"
."
.~
.,
'5
.!!i
.
~
0
~
~
~
§v
.ev
:~
x
w
~
,
0.
0
c,
i;
E
0
e~
0
v
2
."
0
C
'"
.~
."
:".
.~
.;;"
0
3
.g
.0
.a
'"
-,~,
.
"., ~.
~ Intrinsic disorders associated with muscle damage. Clinical examples. *
*This flow chart is designed for classroom or group discussion. Flow chart is not meant to be exhaustive.
Biomechanics
of Joints
Biomechanics
of the Knee
Margareta Nordin, Victor H. Franke
Introduction
Kinematics
Range of Motion
Surface Joint ~J1otion
Tibiofemoral Joint
Patellofemoral Joint
Kinetics
Statics of the Tibiofemoral Joint
Dynamics of the Tibiofemoral Joint
Stability of the Knee JOint
Function of the Patella
Statics and Dynamics of the Patellofemoral Joint
Summary
References
Introduction
The knee transmits loads, part III pales in motion,
aids in conservation of momentum, and provides a
force couple for activities involving the leg. The human knee, the largest and perhaps 1110st complex
joint in the bod)!, is a two-joint structure composed
'of the tibiofemoral joint and the patellofcmoral
joint (Fig. 7-1). The knee sustains high forces and
moments and is situated between the body's two
longest lever arms (the femur and the tibia), making
it particularly' susceptible to injury. This d18ptcr utilizes the knee to introduce the basic terms, explain
the methods, and demonstrate the calculations necessary for analyzing joint motion and the forces and
moments acting on a joint. This information is np·
plied to other joints in subsequcnt chaptcrs.
The knce is particularl.. . . \vell suited for dem
strating biomechanical anal.vscs of joints beca
these anal.. .,scs can be simplified in the knee and
yield useful data. Although knee motion occurs
multancous!y" in three planes, thc motion in
planc is so great that it accounts for nearly all of
motion. Also, although many muscles prod
forces on the knee, at any particular instant
muscle group predominates, generating a force
great that it accounts for most of the muscle fo
acting on the knee. Thus, basic biomedwni
analyscs can be limited to motion in one plane
to the force produced by a single muscle group
still give an understanding of kncc mc)tion and
estimation of the magnitude of the principal for
and moments on the knee. Advanced bioIllcchan
d.\'namic anal,\'ses of thc knee joint that include
Anterior cruclate
ligament
Medial
collateral
ligament
Popliteus
tendon
Fibular
collateral
ligament
Semimembrano
Biceps
femoris
tendon
Superficial
medial collat
ligament (cu
Iliotibial
band (cut)
Interosseous
membrane
Transverse
ligament
-cr--l'-
Gracillis
Semitendinosus
Patellar tendon
B
Two-joint structure of the knee. A, Lateral view of a knee joint with open growth plates.
B, Anterior view without patella.
Sartorius
soft tissue structures are complex and still under investigation.
Analysis or motion in any joint requires the lise of
kinematic daw. Kinematics is the branch of mccrlanics thal deals with Illotion of a body without
reference to force or mass. Anal:vsis of the forces
and momenl~ acting on a joint necessitates the
usc of both kinematic and kinetic data. Kinetics is
the branch of mechanics that deals with the motion
of a body under the action of given forces and/or
moments.
Kinematics
Kinematics defines thl..' range of motion an
scribes the sw-facc motion of a joint in three p
frontal (coronal or longitudinal), sagittal, and
verse (horizonlal) (Fig. 7 2, A & B). Clinical
surements of joint range 01" motion defin
anatomical position as a zero position for mc
ment. This laxonomy will be uscd for joint m
throughout this book. Other taxonomies and
ence systems exist (Ant!riacchi et aI., 1979; Gr
H
Proximal/
dislal
!j"~~;;:~i..k~~~i===::::=Fronlal
~1
plane
•.,
I~.
o
k'.
~
-Sagittal
plane
;,.j
)1
~
Internal/external
Varus!
valgus
A
A, Frontal (coronal or longitudinal), sagittal. and transverse
(horizontal) planes in the human body performed easily for
both the tibiofemoral and the patellofemoral joint. B, Depiction and nomenclature of the six degrees of freedom of knee
motion: anterior posterior translation, medial/lateral transla-
.,. . .<:,
";'~'
,
I?:A
B
tion and proximal distal translation, flexion-extension
tion, internal-external rotation, varus-valgus rotation.
Adapted from Wilson, S.A., Vigorita, V.J., & Scott, W.N
(1994). Anacomy. In N. Scott (Ed.), The Knee (p. 17). Philad
Mosby- Year Book.
Sun tal' 1983; Kroemer ct aL. 1990; bzkaya &
Nordin, 1999), but the anatomical reference system
bv far is the mOSl commonlv lIsed among clinicians.
the two joints composing the kne'; the tibiofemoral joint lends itself particularly well to an
analysis of range orjoinl motion. Analysis of surface
joint motion can be performed easily For both the
tibiofemoral and the P<'11cllofcmoral joint. An~' impediment of range of motion or surface joint 'motion
will disturb the normal loading pattern of' a joint
a.nd bear consequences.
O'r
-20 !:=------;!':---7::----==---:'::-100
20
40
60
80
10
RANGE OF MOTION
The range
Percentage of Cycle
or motion of any joint can be measured in
any plane. Gross measurements can be made with a
goniometer, but more specific measurements require the lise or more precise methods such as elcclrogoniomelr~·. roenlgenogJ"aphy. stcreopllologralllmetr)'. or photographic and video techniques using
skeletal pins,
In the tibiofcmoral jail'll. nlOtion takes place in all
three planes, but the range of motion is greatest by
far in the sagittal plane. Motion in this plane from
full extcnsion to Fulillexion of the knee is from 0' to
approximately 140°.
Motion in the transverse plane. intel-nnl and external rotation. is inOuenced by the position of the
joint in the sagittal plane, With the knee in Full extension. rOHHion is almost completely restl'ictcd b~f
the interlocking of the femoral and tibial condyles.
which occurs mainly because the medial femoral
condyle is longer than the lateral condyle. The range
of rotation increases as the knee is flexed, reaching
a maximum at 90° of nexion; with the knee in this
:- position. external rotation ranges from 0° to approximately 45° and internal rotation ranges from
0' to appmximatcly 3D'. Beyond 90' of Ilexion. the
range of internal and external rotation decreases,
primarily because the soft tissues restrict rOlation.
Motion in the frontal plane. abduction and adduction. is similarly aFfected by the amount of joint
flexion, Full extension of the knee precludes almost
all motion in the frontal plane. Passive abduction
and adduction increase with knee Oexion up to 30°.
but each reaches a maximum of only a few degrees.
With the knee flexed beyond 3D'. motion in the
frontal plane again c1ecre~lses because of the limiting function of the sol't tissues.
The range of tibiofernoral joint motion required
For the performance of various physical activities
can be determined from kinematic anal~lsis. MOlion
Range of motion of the tibiofemoral joint in the sagittal
plane during level walking in one gait cycle. The shaded
dred indicates variation among 60 subjects (age range 20
to 65 years). Adapted {rom MtJ((a}~ M.P., Drought, A.B., Kor}
R. C. (1964j. I/o/alking patterns of norma! men. J Bone Joint Su
46A,335,
in this joint during walking has been measured
all planes. The range of motion in the sagittal pla
during level walking was measured with an e1ectr
goniometer by Lamoreaux (1971) and MUlTay et
(1964). Full or nearly Full extension was noted at t
beginning of the stance phase (0% of cycle) at he
strike, and at the end of the stance phase before LO
off (around 60% of cycle) (Fig, 7-3), Maximum fle
ion (approximately 60°) was observed during t
middle of the swing phase (sec Chapter 18, Biom
chanics of Gait. For more detailed inFormation
These measurements are velocit.y-c1ependent an
must be interpreted with caution.
~"lotion in the transverse plane during walking h
been measured by several investigators. Using a ph
lOgraphic technique involving the placement
skeletal pins through the femur and tibia, Leve
and associates (1948) Found that total rotation of t
tibia with respect to the fernur ranged from appro
imately 4 to 13' in 12 subjects (mean 8,6'). Great
rotation (mean 133') was noted by Kettclkamp an
coworkers (1970). who used electrogoniollletry o
22 subjects. In both studies, external rotation beg
during knee extension in the stance phase an
reached a peak value at the end of th~ swing pha
just before heel strike. Internal rotation was note
during Ilexion in the swing: phase.
DDIII-~
Range of Tiblofemoral Joint Motion In the
Sagittal Plane During Common Activities
I
Activity
Range of Motion
from Knee Extension
to Knee Flexion (Degrees)
- - - - - - - - -0-67"
------,
Walking
""I
Climbing stairs
Descending stairs
Sitting down
Tying a shoe
lifting an object
0-90
0-93
0-106
0-117
';Dillil lrom Kel1clkamp el al. (1970i. fvlean for 22 subjects. A slight
ddiercncc was found bel'.'.'een r19ht "nc! lei! knees (mean for right
knee 68.1°, nW,ln for left knee 66,r).
"These and sub:.equerH date' from lauberllhal Cl al. (1972). Mean for
30 subjects.
Kcltclkamp's group (1970) also mcasurcd motion
in the frontal plane during walking. In nearly all of
the 22 subjects, maximal abduction of the tibia was
observed during extension at heel slrike and at the
beginning of the stance phase; maximal adduction
occurred as the knee \Vas flexed during the swing
phase. The total amount of abduction and adduction averaged 11°.
Values for the rangc of motion o!" the tibiofemoral
joint in the sagittal plane during several comn101l
activities arc presented in Table 7-1. l'vlaximal knee
flexion occurs during lifting. A range of motion
I'rol11 full extcnsion to at least II r 01' flcxion appears to be required for an individual to carry out
the activities of daily living in a normal manner. Any
restriction of knee motion can be compensated for
by increased motion in other joints. In studying the
range of libiofcmoral joint motion during various
activities, researchers found that an increased speed
of movement requires a greater range of motion in
thc tibiol'cl11oral joint (Holdcn et aI., 1997; Pcrry ct
aI., 1977). As the pace accelerates from walking
slowly to running. progressively more knee flexion
is nccdcd during thc stancc phasc (Table 7·2),
SURFACE JOINT MOTION
SUrract~
joint motion. which is the Illation between
the articulating surfaces of a joint. can be described
1'01' any joint in any plane with the lise of stcrcophoLogrammetric mcthods (Sclvik, 1978, 1983)~ Bc~
callsc lhese mcthods ~lre highly technical and com
plex, a simpler method evolved in the nineteen
ccnL'It'\, Is sLillus~d (Reuleau,\, 1876), This m~,h()
called the installt cl.?nter techniqul.', allows surfa
joinl 1l1Olion lO be analyzed in the sagiual an
froolal planes but not in the transverse plane. T
instant center l~chniqLle pl{)vides a description
the rch'tive uniplanar motion of two adjacent se
ments or n body and the direction of displaceme
or the contact points bct\\'L~cn these segmcnts,
The skeldal portion 01" a body segment is called
link. As one link rotaLes ~il){H1l the other. al any i
stum then... is a point that docs nOl mov(:, that is,
point thal has h'ro velocil.V. This point conslitlli
an instantaneous cenLer of motion, or instant ce
tel: The instant center is found b.\' identirying t
displacen1cnt (If two points on a link as the li
moves from one position to another in relation to
adjacl.?nt link, which is considered to he stalionar
The points on thl.' moving link in its original po
tion and in its displaced position arc designated
a graph and lines an.' drawn connecling the two se
of points. The pCl-pcndicular bisectors or these tw
lines are then drawn. The intersection or the pe
pendicular bisectors is thc instant ccntcl:
Clinir::'lll~', a path\\'a~' of the instant center for
joint call be determined b\' takinQ
- successive roen
genograms of the joint in dilTcrent positions (usual
10° apart) throughout the range of motion in o
plane and applying the Rculcaux melhod for locali
the inslanl center 1'01' each intcrv[ll or motion.
\Vhcn the instant center palhway has been dete
mined 1'01' joint motion in one plane, the surfa
joint mol ion C[ln be described, For each intel'val
motion, the point at which the joint surfaces ma
contact is located on the roentgenograms lIsed f
the instant center analysis, and a line is drawn frol
~
~
1Imtl!JIJI--~
I
-~~-,-~----~
Amount of Knee Flexion During Stance
Phase of Walking and Running
Activity
Range in Amount of Knee Flexio
During Stance Phase (Degrees)
W(llking
Slow
Free
Fast
Running
Data
0-6
6-12
12-18
18-30
itom Perry ct al. (1977). Range for seven subjects.
the instant center [() the contact point. A second line
drawn at right angles to this line indicates the direclion of displacement of the contact points. The
direction of displacement of these points throughout the range of motion describes the surface motion in the joint. In most joints, the instant centers
lie at a distance from the joint surface, and the line
indicating the direction of c1isplaccrncnt o!" the conwct points is tangential to the load-bearing surface,
demonstrating thm one joint surFace is gliding on
the other (load-bcaring) surface. In the casc in
which the instant center is found on the sudace. the
joint has a rolling motion and there is no gliding
function. Because the instant center technique allows a description of motion in one plane only, it is
not useful for describing the surface joint motion if
more than I5° of motion takes place in any plane
other than the one being rneasured.
In the knee, surface joint motion occurs between
the tibial and fcmoral condyles and between the
fcmoral condylcs and the patella. In thc tibiofemoral
joint, surface motion takes place in all three planes
simultaneously but is considerably less in the tra
verse and frontal planes. Surface motion in
patellofcmoraljoint takes place in two planes sim
tancoLisly, the frontal and transverse, but is
greater in the frontal plane.
Tibiofemoral Joint
An example will illustrate the lise of the instant c
ter tcchniquc to describe the surface motion of
tibiofemoraljoint in the sagittal plane. To determ
the pathway of the instant center of this joint dur
flexion, a lateral roentgenogram is taken of the k
in full extension and successive films arc Laken
10° intervals of increased nexion. Care is taken
kcep the tibia parallel to the ,-ray table and to p
vent rotation about the femur. \Vhcn a patient
limited knee motion, the knee is Ilc,ed or e'tend
only as far as the patient can tolerate.
Two points on the femur that arc easily iden
ned on all roentgenograms are selecled and des
Locating the instant center. A, Two easily identifiable points on the femur are designated
on a roentgenogram of a knee flexed 80°. B. This roentgenogram is compared with a
roentgenogram of the knee flexed 90°, on which the same two points have been indi~
cated. The images of the tibiae are superimposed, and lines are drawn connecting each set
of points. The perpendicular bisectors of these two lines are then drawn. The point at
which these perpendicular bisectors intersect is the instant center of the tibiofemoral joint .
for the motion between 80 and 90° of flexion. Courtesy of Ian Goldie, M.D. Univ~rsity of
Gothenburg, Gothenburg, Sweden.
natcd on each roentgenogram (Fig. 7.4..4). The
films are then compared in pairs, with the images
of the tibiae supcl-imposed on each othcl:
Roentgenograms with marked differences in tibial
alignrncnt arc not used. Lines are drawn between
the points on thc felllur in the two positions, and
the perpendicular bisectors of these lines arc then
drawn. The point at which these perpendicular bisectors intersect is the instant center or the
tibiofemoral joint for each 10° interval of motion
(Fig. 7-4B). The instant center pathway throughout
the entire range or knee Oexion and extension can
then be ploued. In a normal knee, the instant center pathway for the tibiofemoral joint is semicircular (Fig. 7-5).
After the instant center padnvay has been determined for the tibiofemoral joint. the surface
motion can be described. On each set of superimposed roent-gcnograms the point of contact of the
tibiorcmoral joint surfaces (the narrowest point in
the joint space) is determined and a line is drawn
connecting this point with the instanl centet: A
second line drawn at right angles to this line inclicates the direction of displacement of the contact
points. In a normal knee, lhis line is tangential to
the surFace of the tibia for each interval of motion
from full extension to full flexion, demonstrating
that the femur is gliding on the tibial condyles
(Frankel et aI., 1971) (Fig. 7-6). During normal
knee motion in the sagittal plane from full extension to full flexion. the instant center pathway
moves posteriody, forcing a combination of
rolling and sliding to occur between the articular
surface (Fig 7-6. A & B). The unique mechanism
prevents the femur from rolling off the posterior
aspect of the tibia plateau as the knee goes into increased flexion (Draganich et aI., 1987; Fu et aI.,
1994; Kapandji, 1970). The mechanism that prevents this roll-off is the link formed between the
tibial and Femoral attachment sites of the anterior
and posterior cruciate ligaments and the osseous
geometry of the femoral condyles (Fu el aI., 1994)
(Fig. 7-6, B-D).
Frankel and associates (1971) determined the instant center pathway and analyzed the surface motion of the tibiofemoral joint from 90· of Ilexion to
Full extension in 25 normal knees; tangential gliding was noted in all cases. They also determined the
instant center pathway for the tibioFemoral joint in
30 knees with internal derangement and found
that, in all cases, the instant center was displaced
from the normal position during some portion of
the motion examined. The abnormal instant cen-
Semicircular instant center pathway for the tibiofemor
joint in a 19-year-old man with a normal knee.
tel' pathway for one subject, a 35-~"ear-old m
with a bucket-handle derangement. is shown
Figure 7-7.
If the knee is extended and Hexed aboul an
normal instant center pathway, the libiofem
joint surfaces do not glide tangentially throu
out the range of motion but become cither
tracted or compressed (Fig. 7-8). Such a kne
annlogous to a door with a bent hinge that
longer fits into the door jarnb. If the knee is c
tinually forced to move about a displaced ins
center, a gradual adjustment to the situation
be reflected either b~: stretching of the ligam
and other supporting soft tissues or by the im
sition of abnormall~' high pressure on the art
lar surfaces.
Internal derangements of the libiofcmoralj
may interfere with the so-called screw-h
mechanism, which is external rotation during
tension of the tibia (Fig. 7-9). The tibiofem
~ioint is not a simple hinge joint; il has a spiral
helicoid, motion. The spiral motion or the t
about the femur during flexion and extension
sults from the anatomical configuration of
medial femoral condyle; in a normal knee, th
condyle is approximately 1.7 ern longer than th
lateral condyle. As the tibia moves on the femu
from the fully flexed to the fully extended pos
tion, it descends and then ascends the curves
the n1edial femoral condyle and simultaneous
rotates externally!. This motion is reversed as th
tibia moves back into the fully flexed positio
This screw-home mechanism (rotation around th
longitudinal axis of the tibia) provides more st
bility' to the knee in any position than would
simple hinge configuration of the tibiofemor
joint.
Matsumoto et al. (2000) investigated the ax
of tibia axial rotation and its change with kne
flexion angle in 24 fresh-frozen normal knee ca
daver specimens ranging in age From 22 to 6
years. The magnitude and location of the long
tudinal axis of tibia rotation were measured
15° increments between 0 and 90° of knee fle
ion. The magnitude of tibia rotation was 8° at
Gliding
A
8
c
I
D
1mIlI'------I
_
A. In a normal knee, a line drawn from the instant center of
the tibiofemoral joint to the tibiofemoral contact point
(line A) forms a right angle with a line tangential to the tibial surface (line B). The arrow indicates the direction of displacement of the contact points. Line B is tangential to the
tibial surface, indicating that the femur glides on the tibial
condyles during the measured interval of motion. B. Pure
sliding of the femur on the tibia with knee extension. Note
that the contact point of the tibia does not change as the
femur slides over it. Eventually impingement would occur if
all surface motion was restricted to sliding. Round points
delineate contact points at the femur and triangles delineate contact points at the tibia. C. Pure rolling of the femur
on the tibia with knee flexion. Note that both the tibia and
the femoral contact points change as the femur rolls on the
tibia. Also note that with moderate flexion, the femur will be~
gin to roll off the tibia if surface motion was restricted to rolling. D, Actual knee motion including both sliding and rolling.
Abnormal instant center pathway for a 35-year-old man
with a bucket-handle derangement. The instant center
jumps at full extension of the knee. Adapted from Frankel,
VH., Burstein, A.H., &- Brooks, D.B. (1971). Biomechanics of in
ternal derangement of the knee, Parhomechanics as determin
by analysis of the instant centers of motion. J Bone Joint Surg,
53A, 945.
of knee flexion. The tibial rotation increased
rapidly as the knee flexion angle increased and
reached a ma:dmunl of 31 0 at 30 0 of knee flexion.
IL then decreased again with additional flexion
(Fig 7-10). The location of the longitudinal rotational axis was close to the insenion of the anterior cruciate ligament (ACL) at 0° of flexion. At
continuous flexion up to 60°, the I'otational axis
moved toward the insertion of the posterior cruciale ligament. Belween 60 and 90° of flexion,
the rotational axis moved anteriorly again (Fig
7-11). This study showcd thaI lhc rolational axis
remains approximately in the area between the
two cruciate ligaments. Any change of direction
and tension of the cruciate ligaments and surrounding soft tissue may affect the movement
and the 'location of the longitudinal tibia axis
of rotation and thcrcby affccl joint load distri-
~-L'-JI
Extension
\'1
'.,
... External
rotalion during
extension
Screw·home mechanism of the tibiofemoral joint. D
knee extension. the tibia rotates externally. This mo
reversed as the knee is flexed. A. Oblique view of t
mur and tibia. The shaded area indicates the tibial
solid line axis for knee flexion and extension, dotte
internal and external rotation axis of the tibia duri
ion and extension. AcJapred from Helier. A.J. (1974j. A
and mechanics of movement of rite knee Joinr. In A. He
Disorders of the Knee fpp. i -17). Philadelphia.- J B. Lipp
bution.
A clinical tcst. the Helfet test, is often useel to determine whether external rotation of the tibio-
•
Distraction
A
Compression
B
Surface motion in two tibiofemoral joints with displaced
instant centers. In both joints, the arrowed line at right an·
gles to the line between the instant center and the
tibiofemoral contact point indicates the direction of displacement of the contact points. A. The small arrow indicates that with further fIeldon. the tibiofemoral joint will
be distracted. B, With increased flexion, this joint will be
compressed.
femoral joint takes place during knee eXl
thereby indicating whether the screw-home m
nism is illlact. This clinical tcst is performe
the patient sitting with the kn~e and hip flex
ancl the leg hanging free. The medial and
borders of the patella are marked on the ski
tibial tuberosity and the midline of the pate
then designated, and the alignment of the
tuberosity with the patella is checked. In a n
knee Hexed 90°, the tibial tuberosily aligns w
medial half or the patcll" (Fig. 7-121\). The k
then extended fully and the movement of the
tuberosity is obsel\:cd. [n a normal knee. the
tuberosity moves laterally during extensio
"Iigns with the latent! h"lf or the patella at
tension (Fig. 7-128). RotatOl')' motion in a n
knee may be as great as hair the width
patella. In a deranged knee, the tibia may not
externally during extension. Because of the
surface motion in such a knee, the tibiof
joint will be abnormally compressed if the k
forced into extension, and the joint surfaces m
d.amaged.
patellofemoral Joint
The surface motion of the patellofcmoral joint in
the frontal plane may also be described by means or
the instant center technique. This joint is shown to
have a gliding motion (Fig. 7-13). From full extension to full flexion of the knee, the patella glides
caudally! approximately 7 em on the femoral
condyles. Both the medial and lateral facets of the
femur articulate with the patella from Full extension to 140' of flexion (Helme, 1990) (Fig. 7-14). Beyond 90' of flexion, the patella rotates externally,
:md only the medial femoral facet articulates with
the patella (Fig. 7-14B). At full flexion, the patella
sinks into the intercondylar groove (Goodfellow et
aI., 1976). The contact area of the lateral facet joint
of the patella is larger than the medial contact areas and ranges from 0.5 to 2.5 cn1~ and less than 0.5
to 2 em"' ) respectively" Contact areas increase \vith
an increased amount of flexion of the knee joint
and increased pulling force of the quadriceps muscle (I-lehne, 1990).
-r
-r
-.§
ro
30
-0
a:
ro
·x 20
-- --
r-~
r-~
-
--
--
.............
Tibia Plateau
"
.. ..
'
..........
ACL
•
0·'0
\.
5'
Lateral
'.
::.....
'. ....
Fibula
30
". •••••••• ..45"
6t?~L,.
.
o
• •••••••
.
Axis locallon an
standard deviat
location of the axis of tibia axial rotation. The location
the axis was close to the tibial insertion of the anterior
cruciate ligament (ACL) at A" of flexion and gradually
moved toward that of the posterior cruciate ligament a
45 and 90" of knee flexion. The axis then moved anteri
again and was approximately at equal distance from th
two insertions of the cruciate ligaments at 90" of knee
flexion. ACL, insertion of the anterior cruciate ligament
PCl, tibial insertion of the posterior cruciate ligament.
Reprinted v'/ith permission from MatsurnolO, H., Seedhom, B
Sue/a. Y, et £11. (2000). Axis of tibial rotation and its change
flexion angle. Clin Orthop, 371, J78-182.
'r"o
:0
i=
ro 10
-0
>-
!
o o
IL -
Kinetics
15
30
45
60
75
.
90
Knee Flexion Angle
jIBm'------The total tibial axial rotation (y-axis) plotted against the
magnitude of knee flexion (x-axis) in fresh-frozen specimens tested under unloaded axial conditions, The magnitude of tibial rotation was below 9" at A" of knee flexion.
Maximum rotation (31.]0) was measured between 30 to 45"
of knee flexion. Reprinted with permission from Matsumoto,
H., Seedhom, B.B., Suda, Y, et al. (2000), Axis of tibial rotation
and its change vi/ith flexion angle. Clin Orlhop, 371, 778-182.
Kinetics involves both static and d,vnamic ana
sis of the forces and moments acting on a jo
Statics is the stud~/ of the forces and moments a
ing on a body in equilibrium (a body' at rest
moving at a constant speed). For a bod)' to be
equilibrium. two equilibrium conditions must
met: force (translator:v) equilibrium, in which
sum of the forces is zero, and moment (rotato
equilibrium, in which the sum of the moment
zero. Dynamics is the study of the moments a
forces acting on a body in motion (an accelerat
or decelerating body), In this case, .the forces
not add up to zero, and the body displaces and
the moments do not add up to zero and the bo
Knee flexed 90c
Knee lully extended
Helfet test. A, In a normal knee flexed 90°, the tibial tuberosity aligns with the medial half
of the patella. B, When the knee is fUlly extended. the tibial tuberosity aligns with the lat·
eral half of the patella .
•
After the instant center (lQ is determined for the
patellofemoral joint for the~ motion from 75 to 90° of knee
flexion, a line is drawn from the instant center to the con·
tact point (CP) between the patella and the femoral
condyle. A line drawn at right angles to this line is tangential to the surface of the patella, indicating gliding .
•
Superior
Lateral
M
Medial
Lateral
.,~
~ '\
(U
-
-
-
-30'-90"
I
I
Inferior
20' 45' 90'
Superior
,
---120"
Lateral
Me
I
Inferior
135'
;0..
A
IiR'm
A, The position of the patella at different ranges of knee
I
flexion motion. a, Contact areas during different degrees of
flexion. Beyond 900 of flexion, the patella rotates slightly
outwards. Adapted from GoodfelJo'J'l, I. Hvngerford, D.5., & lin·
del, M (1976), PateJlofemoral joint mechanics and pathology. I.
B
Functional anatomy of the patellofemaral joint. J 80ne Joint
58B. 287; and from Helme. H.I (1990). Biomechanics of the
parellofemoral joint and its clinical relevance. Clin Onhop. 25
73-85,
•
rotates around an axis perpendicular to the plane
of the fOl'ces producing the moments. Kinetic
analysis allows one to determine the magnitude of
the moments and forces on a joint produced by
body weight, muscle action, soft tissue resistance,
and externally applied loads in any situation, either static or dynamic, and to identify those situations that produce excessively high moments or
forces.
In this and subsequent chapters. the discussi
stalics and dynamics of the joints of the skelela
tem concerns the magnitude of the forces and
ments acting to move a joint about an axis
maintain its position. It does not lake into acc
Ihe deforming effect of Ihese forces and mom
on the joint structures. This effect is indeed pre
but the discussion is not within the s~ope of this
STATICS OF THE TIBIOFEMORAL JOINT
Static analysis may be lIsed to determine the forces
and moments acling on a joint when no motion
takes place or at one instant in time during a dynamic activity such n5 walking, running, or lifting
an objecl. 1t can be performed for any joint in any
position and under any loading configuration. In
such analyses, eHher graphic or mathematical
methods may be lIsed to solve for the unknown
forces or moments.
A complete static analysis involving all moments
and all forces imposed on a joint in three dimensions
is complicated. For this reason. a simplified technique is oflen lIsed. The technique utilizes a frecbody diagrarn and limits the analysis lo one plane, to
the three main coplanar forces acting on the frccbody. and ro the main moments ~cting about the
joint under consideration. The minimum magnitudes of the forces and moments arc obtained.
When the simplified Free-body technique is used
to analyze coplanar forces, one portion of the body
is considered as distinct rrom the entire bod.\', and
all forces acting on this free-body arc identified. A
diagram is drawn of the free-body in the loading situation (0 be analyzed. The three principal coplanar
forces acting on the free-body are identified and
designaLcd on the free-body diagram.
These forces afC designated as vectors if four
characteristics are kno\vn: magnitude, sense. line of
application, and point of application. (The tcrm '"direction" includes line of application and sense.) If
the points of application for allthrec forces and thc
directions for two forces arc kno\vn, all remaining
characteristics can be obtained for a force equilibrium situation. \Vhen the free-body is in equilibrium, the three pl-incipal coplanar forces are concurrent; that is. they intersect at a common point. In
other words, these forces form a closed system with
no resultant (i.e., their vector sum is zero). For this
reason. the line of application for one force can be
determined if the lines of application for the other
two forces arc known. Once the lines of application
for all IIwee forces are known, a triangle of forces
can be constructed and the magnitudes of all three
forces can be scaled from this triangle.
An example will illustrate the application of the
simplified Ii'ee-body techniquc for coplanar forccs to
the knee, In this case, the technique is lIsed to estimate the minimum magnitude of the joint reaction
force acting on the tibiorcmoral joint of the weightbearing leg when the other leg is lifted during stair
climbing. The lower leg is considered as a free-body.
distinct from the rest of the bod.". <'lIlt! a diagram o
this free-bod." in tilt..' stair-climbing situation is
drawn (Calculation Bo,,,; 7-1). I":'nlrn all rorces acting
on thl..: I"rct..··body. the three main coplanal- I"o('(,:(:s an
idenlified a~ the ground rL'action force (equal to
body weight I), the tensile force through the patella
tcndon exel"ted h.v llw quadriceps Illust..:ll' , and th\.
joint reaction force on the tibial plateau. The ground
n.:action force (\rV) has a known m<'lgnitude (L'qu;;llto
body wt..'iglu). S~IlS('. line or <'lpplicalion. and poinl o
application (poii'll or contact between the foot and
the ground). The palellnr tendol1 force (P) has a
known SC'IlSL' (awa.'; from the knee joint), lint..' of ap
plicaliun (along lhe patellar t('ndon). and point
application (point or insertion of thc patellar tendol
on till' tibial tulx.·rosit."), but an unknown l1lagni
tude. The joint n:action force (J) has a known poilU
of application OJl the slIrl"acL' or the tibia (the COJ1lac
point of tilL' joint surfaces betw(.'en the tibial and
femora! cond,vlc:s, estimated from a roelllgenogram
or the joint in the pn)!x~r loading configuration), bu
an unknown magnitude, sense. and line of applica
tion. Using \'L'ctors cakuh\lions <:\11('\ triangles law
the joint rL'action force (.1) and thl' pall.'lIar tendon
force (P) call be c,dculated (Calcul"tion Bo\ 7-1)
It call bt..· seen thaI the main musck forct..' has
llluch gre<.ltL'1' inlhICIll:L' on the magnitude of" th
.joint reaction force than dot..·s the -~round !"(.'actio
force produced b." bod~· \vcight. Note lhat, in this ex
amplc (Calculation Box 7·1), only til(' minimum
magnitude of the joint rt..'~lction force has been
calculated. If other muscle I"oret-'s arc considered
such as tht..' foret..' produced b~' tht..' contraction of th
halllslring Inusc!es in stabilizing the knee, the jClin
reaction force inCl·cast..·s_
Th\.' ne:'.;t Slt.·p in the static analysis is ~\I1;:d."sis o
the moments acting around the centcI- of motion 0
the libiofcrnoral joint with the knee in the same po
sition and lht..' loading. configuralion shown in Ca
culation Bo:'.; Figure 7·1-1. The 1ll01llL'11I <'lIlal.\·sis i
Llsed to cstimak' the minimulll ll1agnitudL' of th
mOIl1L'nt produced through the patellar h:ndon
which counterbalances the 1l10ll1Cnt on the lowcr le
produced b.'; the weight or the hod." as the subjec
ascends stairs (Calculation Box 7·2).
o
C:lSl', the ground I"l'~ll"lillll rurn' is ;u.::lu;t1ly L'qllallo hod
mintl:- (he \\"..::i!..:hl of till' IO\\"L'r k·t:. BL'l";III:-l' tit...· \widll o
the 'it'Hwr k~ is lllillilT~;ll {1L'SS (hall Olli.' 1:'l1(h or Ihl' bOtl\-), i( CI
lin
lhi:,
\\"L'idlt
l.k' di,.. rl·~aranl. ;llld the li~url' for tlll:tl hod\- \\"l'i!.!hl l\\·U bl' tu
lizL'd ill Illl' l·akul:lIion. .
.-
~
!, Free-Body Diagram of the Knee Joint
Force P ..,
~
I
,..
\.\
.<
Known: Point of applicalion
;':"
\Q',
Known:
• ~
Sense
The three main coplanar forces acting on the lower leg:
... Force J
Line of application
Point of application
Unknown: Magnitude
UnknO'.·',n: Magnitude
Sense
\
Ground reaction iorce (W). patellar tendon force (P), and joint
reaction force (J) afe designaled on a free-body diagram of th
lower leg while climbing stairs (Calculation Box Fig 7-1·1),
Line of application
\\
.,•..,-..__ ....\
". -'. \.
<\.
',-.\
I
'.
\
"-
:i~
/\J
o:__ ~~
I
·
I---==~-r"'iKnown:
~-
/'
Magnilude
Force W
Sense
Line of application
!Point of application
.
Because the 100\'cr leg is in equilibriulll, the lines of applicatio
.' Force J
" "'i
v~
.:.~
for all three forces intersect at one point. Because the linesof
l~~:\
T ibialemoral
Force P \ ( \ ..... i
contact point
\~ \.
,,
[I,.VO
forces (Wand P) are known, the line of a
plication for the third force (;) can be determined. The lines o
application for forces Wand P are extended until they jnter~
'.'.
.\:;....
sect. The line or application for J Gill then be drawn from its
,
\\........
'.
\
point of application on the tibial surface through the intersec
\
tion point (Calculation Box Fig.
Intersection
~\
pOint~
I
application for
7-1-2).
-----·-·-~·=::::·:"~~W-1J.-·'""'---~----
Now that the line of application for J has been determined, it
is possible
\'.,,
I
Force J
4.1W
Force P
,, 3.2W
. ,,
.... \,
'. ,
'. ,
'. ,
'. ,
'. ,
'.
'.
'.
'.
_
drawn from the head of ve([or W. Then, to close the triangle
force J is drawn from rhe head of vector W. The point at whic
Now that the length of all three vectors is known, the magni
tude of forces P and J can be scaled from force W, which is
. . .J
.
constru([ a triangle of forces (Calculation Box F
forces P and J intersect defines the length of these vectors.
Force W
{,
(0
7-1 ~3). First, a ve([or representing W is drawn. Next. P is
equal to body weight. It is determining the number of limes
the length of force W can be aligned along the' force P and J,
respectively. In this case, force P is 3.2 times body weight, and
force J is 4.1 times body weight.
Free-Body Diagram of the Lower Leg During Stair Climbing
The
two main mornents acting around tt1e center of motion
the tibioiemoral JOlfH (solid dOl) are designated on the free
body diagram of the lower leg during stair climbing (Calcul
(
tion Box Fig. 7-2-1).
The fleXing moment on the lower leg is the product of t
\\
weIght of the body:- (W, the ground reaction force) and its
\/\
\
fever arm (a), 'Nhich is the perpendicular distance of the for
Force P \
W to the center 0; motion aT the tibiofemoral joint. The co
terbalanCIng extending moment IS the product of the quad
ceps muscle force through the patellar tendon (P) and its le
arm (b), Because tile lo""er leg is in equilibrium.
the sum
of
these two moments must equal zero (~M = 0).
In this example. the counterclockwise moment is arbitrari
designated
as positive (W x a - P ~< b
= 0). Values for lever
arms a and b can be measured from anatomical specimens o
on soft tissue imaging or fluoroscopy (Kellis & Baltzopoulos.
1999; Wretenberg et aI., 1996), and the magnitude of W ca
be determined from the body weigt1t of the individual. The
magnitude of P can then bE' found from the moment equilib
:::M = 0
num equation:
W;.:
a
p~_.-
b
W>:a=Pxb
p =W
!Again the welghl of the lower leg IS disregarded because it is
lESS ,han Oile \Cntil of body vveight
xa
b
I Calculation Box Figure 7-2-1.
e
------..- - - - -
DYNAMICS OF THE TIBIOFEMORAL JOINT
Although estimations of the magnitude of the
forces and moments imposed on a joint in static
situations are llseful. most of Ollr activities are of a
dynamic nature. Analysis of the forces and moments acting on a joint during motion requires the
use of a different technique for solving dynam,ic
problems.
As in static analysis, the main forces considered
in dynamic analysis arc those produced by body
weight, muscles, other soft tissues, and externally
applied loads. Friction forces arc negligible in a normal joint and thus not considered here. In dynamic
analysis. two factors in addition to those in static
analysis must be taken into account: the acc
tion of the body part under consideration an
mass moment of inertia of the body part. (The
moment of inertia is the unit used to expres
amount of torque needed to accelerate a bod
depends on the shape of the body.) (For mo
depth studies of dynamics, sec Ozkaya & No
1999.)
The steps for calculating the minimum m
tudes of the forces acting on a joint at a particu
stant in time during a dynamic aClivity arc as fo
I. The anatomical structures arc identified:
nitions of structures. anatomical landma
point of contact of articular surface. and
2.
3.
4.
5.
6.
:f
4\rl11s involved in the production of forces for
the biol11cchanical analyses.
The angular acceleration or the moving body
part is determined.
The mass monlent of inenia of the moving
bod~" part is determined.
The torque (moment) acting abollt the joint is
calculated.
The magnitude of the main muscle force accelerating the body pan is calculated.
The magnitude of the joint reaction force at a
particular instant in time is calculated by static analysis.
In the first step, the structures of the body involved in producing forces on the joint are identified. These are the moving body part and the main
muscles in that body part that arc involved in the
production of the motion. Great care must be taken
in applying this first step. For example, the lever
arms for all rnajor knee muscles change according
to the degree of knee Ilexion and gender (\Vretenberg et aI., 1996).
In joints of the extremities. acceleration of the
body part involves a change in joint angle. To
determine this angular acceleration of the moving
body part, the entire movement of the body part
is recorded photographically. Recording can be
done with a stroboscopic light and movie camera.
with video photogrummetry, with Selspot systems,
with stereophotogrammctry, or with other methods
(Gardner et aI., 1994; Ramsey & Wretenberg, 1999;
\Vintel~ 1990). The maximal angular acceleration for
a particular motion is calculated.
Next, the mass moment of inertia for the moving
body part is determined. Anthropometric data on
the body parl can be used for this determination.
As calculating these data is a complicated procedure, tables are commonly used (Drillis et aI.,
1964).
The torque about the joint can now be calculated
using Newton's second law of motion, which stales
that when motion is angulat~ the torque is a product
of the mass moment of inenia of the body part and
the angular acceleration of that part:
T = Ie< ,
where
T is the torque expressed in newton meters (Nm)
I is the mass Illoment of inertia expressed in
newton meters x seconds squared ( m sec~)
a is the angular acceleration expressed in radians
per second squared (rlsec').
,
'i; :
.. , .'"
Th..: torque is not only a product or the mass
ment of inertia and the angular acceleration o
body part but also a product or the main Ill
force accelerating the body part and the perpe
ular distance of the force from the center of m
of the joint (level' arm). Thus.
T = Fe!
where
F
d
is the force expressed in newtons (N)
is the perpendicular distance expressed in m
leI'S (01).
Because T is known and d can be measure
the body part from the line of application o
force to the center of motion of the joint, the e
tion can be solved fOl" F. vVhcn F has been c
lated, the remaining problem can be solved l
static problem lIsing the simplified free-body
nique to determine the l11inirnum magnitude o
joint reaction force acting on the joint at a ce
instant in time.
A classic example will illustrate the usc o
narnic analysis in calculating the joint reaction
on the libiofemoral joint at a particular in
during a dynamic activity (c.g., kicking a foo
(Frankel & Burstein, 1970). A stroboscopic fi
the knee and lower leg was taken, and the an
acceleration was found £0 be maximal at the in
the foot struck the ball: the lower leg was almos
tical at this instant. From the film. the maxima
gular acceleration was computed to be 453 !
From anthropometric data tables (Drillis e
1964), the mass momenl of inertia for the lo\v
was determined to be 0.35 Nm sec~. The to
about the tibiofemoral joint was calculated ac
ing to the equation; torque equals mass mome
inertia times angular acceleration (T = (0),
0.35 Nm sec' x 453 rlsec'
=
158.5 Nm
After the torque had been determined to be
Nm and the perpendicular distance fran) the
ject's patellar tendon to the instant center fo
tibiofemoraljoint had been found to be 0.05 m
muscle force aCling on the joint through the pa
tendon was calculated using the equation to
equals force times distance (1' = Fd),
158.5 Nm = F x 0.05 m
F .=
158.5 Nm
-=--=-0.05 m
F = 3170 N
Thus, 3,170 N was the maximal force exerted by
the quadriceps muscle during the kicking motion.
Static analysis can now bt: performed to determine the minimum magnitude of the joint reaction
force 011 the tibiofcrnoraI joint. The main forces on
this joint arc id':lltificd as the patellar tendon force
(P). Ihe gravilaliomtl force of Ihe lower leg
and
the joint reaction force (J). P and T arc known vectors. J has an unknown magnitude. sense, and line
of application. The free-body technique for three
coplanar forces is L1sed to solve for J, which is found
10 be only slighrly lower Ihan P.
As is evident from the calculations, the two main
factors that influence the magnitude of the forces on
a joint in d)'namic situations are the acceleration of
the body part and its mass moment of inertia. An increase in angular acceleration of the body part will
produce a proportional increase in the torque about
the joint. Although in the body the mass moment of
inertia is anatomically set, it can be manipulated extenl~l)Jy. For example. it is incrc.:ascd whcn a weight
boot is applied (0 the foot during rehabilitative exercises of the extensor muscles 01" the knee. Normally, a joint reaction force of approximately 50°;(;
of body weight resulls when the knee is slowly (with
no acceleration forces) extended from 90° of flexion
to full extension. In a person weighing 70 kg, this
force is approximately 350 N. If a IO-kg weight boot
is placed on the foot, it will exert a gravitational
force or 100 N. This will increase Ihe joint reaclion
force by 1,000 N, making this force aimosl rour
times greater than it would be without the bool.
Dynamic analysis has been L1sed to investigate
the peak magnitudes of the joint reaction forces,
muscle forces, and ligament forces on the tibio~
femoral joint during walking. Morrison (1970) calculated the magnitude of the joint reaction force
transmiucd through the tibial plateau in male and
female subjects during level walking. He simultane~
ously recorded muscle activity elcctromyographi~
cally to determine which muscles produced the
peak magnitudes of this force on the tibial plateau
during various stages of the gait cycle (Fig. 7-15).
Just after heel strike, the joint reaction force
ranged from two to three times body weight and
\vas associated with contraction of the Iwmstring
muscles, which have a decelerating and stabilizing
effect on the knee. During knee llexion in the beginning of the stance phase. the joint reaction force
waS approximately lWO rimes body weight and was
associated with the contraction of the quadriceps
muscle, which acts 10 prevent buckling of the knee.
The peak joint reaction force occUlTed during the
en,
..
late s(;Jncc phase just before loe-ofr. This
ranged from two to four times body weight. v
among the subjects tesled, and was <:Issociatcd
contraction of the gastrocnemius muscle. In th
swing phase, contraction of the hamstring m
resulted in a joint reaction force approxim
equal to bod)' weight. No significant dilTcrcnc
found between the joint reaction force magn
for men ancl women when the values were no
izcd by dividing them by body weight.
Andriacchi & Sirickiand (1985) sludied Ihe n
moment patterns around the knee joint during
walking for 29 healthy volunteers (15 women a
men wilh an average age of 39 years). Figure
depicts the flexion-extension, abduction~addu
and internal-external moments during the s
and swing phase of level walking. The momen
normalized (0 the individual's body weigh
height and are presented as a percentage. The f
4
en
I
3
.cen
'Q;
~
'"
'C
0
e
~
~
0
u.
2
fl
./\ ,/\
\/ \! \
.'
/-...,
l
V,..' ' ' 1\
'
\
' ., . .,-.
o1~O:::O-c.._-'-'..J·~---~"'6~O~--'-"'------'~--:1
Percentage of cycle
Joint reaclion
..... Hamslrings
HS
=.
Heel strike
. - -- Quadriceps le
TO
=.:
Toe-ofl
-- .. Gastrocnemiu
Joint reaction forces expressed as body weight trans
through the tibial plateau during walking. one gait
(12 subjects). The muscle forces producing the peak
tudes of this force are also designated .. Adapted from
Morrison, lB. (1970). The mechanics of the knee joinr in
rion lO normal w<lfking. J Biomech, 3, 51.
•
'!'·
extension moments during the stance phnse are approximately 20 to 30 tilTICS larger than the moment
produced in the frontal (abduction-adduction) and
transverse (intcrnal-extcl-nal) planes.
An increase in knee joint flexion-extension mo-
A_
Flexion-Exlension
Moment Pallerns
(Exlernal)
ment amplitude has been reported at increased
walking speeds (Andriacchi & Strickland. 1985;
Holden et aL. 1997). An increase in the production
of adduction knee joint moment during stair climbing compared with level walking was reported b.\' Yu
et al. (1997).
During the gait cycle, the joint reaction force
shifts from the medial to the lateral tibial plateau. In
the stance phase, when the force I'caches its peak
value, it is sustained mainly by the medial plateau
i.
(adduction moment); in the swing phase, when the
force is 111inimal. it is sustained primarily by the lat·
eral plateau, The contact area of the medial tibial
plateau is approximately 50% larger than that of the
lateral tibial plmeau (Keuclkamp & Jacobs, 1972).
Also, the cartilage on this plateau is approximately
three times thicker than that on the lateral plateau.
The larger surface area and the greater thickness of
the medial plateau allow it to more easily sustain
the higher Forces imposed on it.
In a normal knee, joint reaction forces are sustained by the menisci as well as by the articular cartilage. The function of the menisci was investigated
by Seed hom and coworkers (1974), who examined
the distdbution of stresses in knees of human autopsy subjects with and without menisci. Their results suggest that in load-bearing situations, the
magnitude of the stresses on the tibiofemoral joint
when the menisci have been removed may be as
much as three times greater than when these structures are intacL Fukuda et al. (2000) studied in vitro
the load-compressive transmission of the knee joint
and the role of menbci and articular cartilage. The
load simulated was stalic and d).'namic impact loading. The testing was done in neutral. varus, and valgus alignment of the knee joints in 40 fresh-frozen
pig knee speclmens. The compressive stress on the
medial subchondral bone was lip to five times
higher with the menisci removed. This study points
to the importance of the menisci as a structure to
absorb load and protect the canilage and subchondral bone under dynamic conditions.
In a normal human knee, stresses are distributed
over a wide area of lhe tibial plateau. If the menisci
are removed. lhe stresses are no longer disll"ibutcd
Over such a wide area but instead arc limited to a
contact area in the Center of lhe plaleau (Fig. 7-17).
Thus, removal of' the menisci not only increases the
It Pattern 1
¥
>~
;!:
"'
rL
.s0
,
3.0
'xID
u::
1.0
0
I
I
I
I
c
a 1.0
'00
c
2.0
X
'"c w'"
ID
I
I
I
c
a 2.0
'"
3.0
-
B.
x
0
;!:
1.0
c
.~
a
.2
ID
ID
C
"0
"0
,
'"
Abduction-Adduction
Moment Patterns
(External)
1.0
I
"'~
Stance --l-Swing--
13
0
«
2.0
3.0
4.0
5.0
C.
Internal-External
Rotational
Momenl Patterns
(External)
""
I
X
;!:
OJ
1.0
;;;
c
;;; 0.5
E
l,
c
;;;
ID
;;;
x
,
'0
'"c
'"
c
w
0.5
I
I
I
1.0 ..- Stance - .. I'-Swing-"
Flexion-extension (A), abduction-adduction (8), and
internal-external rotation (C) moments produced durin
one gait cycle in normal subjects. The moments are no
malized to each individual body weight X height and
pressed as a percentage. Reprinted with permission from
Andriacch( IP. &. Strickland. A.B. (1985). Gaie analysis as a
to assess joint kinetics. In N. Berme, A.E. Etlgin. O.A. Correis
et al. (Eds.). Biomech(lnics of Normal and Pathological Hum
Articulating Joints. (NATO ASI series. Va/93. 'pp. 83-I02)'
Orodrechr. Netherlands: Marr;nus Nijhoff.
•
•
Force
Force
II I
l
\\
...
III
III
\\\
tIl
\
I
~
tt tt
...
Menisci removed
Normal
Stress distribution in a normal knee and in a knee with the
menisci removed, Removal of the menisci increases the
magnitude of stresses on the cartilage of the tibial plateau
and changes the size and location of the tibiofemoral con·
tact area. With the menisci intact, the contact area encom-
passes nearly the entire surface of the tibial plateau. With
the menisci removed, the contact area is limited to the
center of the tibial plateau.
magniwde of the stresses on the cartilage and subchondral bone at the ceiller of the tibial pl~llcau but
also diminishes the size and changes the location or
the contact area. Over the long term, th~ high
stresses placed on this smaller contact area may be
harmful to the exposed cartilage, which is usually
soft and fibrillated in that area. The menisci are
7.1
thought to carry lip to 70<}(; of the load across
knee. ,Vlovcmcnt during knee Ilcxion of the meni
would therefore protect the articulating surfa
while avoiding injury to it.
Vedi et al. (1999) studied menisci movement in
young football players with normal knees w
MRI. The knee flexion movement was scanned fr
full knee extension lo 90° of knee flexion. The im
ing technique allowed for both standing (weig
bearing) and sitting (non-weight-bearing) and w
performed simultancollsl~t in the sagittal and tra
verse plane. Figure 7-18 shows the movements
the transverse plane of the medial and lateral me
sci expressed in millimeters (mean) from full ext
sion to 90° Ocxion of knee joint motion. Movem
was significantly greater in weight-bearing than
non-\Veigllt~bearing for both lateral and med
menisci. The contributions of the menisci art: lhe
fore not only to protect the articular cartilage a
subchondral bone but also to contribute aClively
knee joint slabilit~:.
STABILITY OF THE KNEE JOINT
The key to a healthy knee joint is joint slability. T
osseous configuration, the menisci, the ligamen
the cnpsulc, and the muscles surrounding the k
joint produce joint slability (Fig. 7-1, A & B). If a
Ant
5.4
Ani
9.5
0
6.3
,
Flexion
3.6
I
3.3r
i
\
N
,
\
\
\
90
"-
f
I
,,
•
-J_3.9
Medial
menisci
"-
Post
\
.,
-L_-
,
"-
5.6
,
3.8
Lateral
menisci
,
-4-
Medial
menisci
Post
.
-~4.0
--
I
,/
Lateral
menisci
B
A
Simplified diagrams showing the mean movements of the
medial and lateral menisci from full knee extension to 90°
knee flexion during two conditions. A. Erect and weight~
bearing. B. Sitting, relaxed. and bearing no weight. Adapted
with permission (rom Vedl, V. WiJjiams. A., Tennant. S.J.. el aJ.
(1999). Menisca! movemenr. An In-vivo srudy u~in9 dynamic M
Bone Join! Swg, 818(1), 37-41.
.ACL Injury
i.
A
J;_
30·year-old male suffered an external rotation trauma
in his right knee while downhill skiing. Following the
~~;,: trauma, he experienced sharp pain, progressive joint effu,~,iqn,_and subjective instability. ,During careful examination
Ii' ,by a specialist, an anterior positive drawer test was dlag-
<";'n,osed, and the Lachman test and pivot shift test were
': found positive. An MRI confirmed the ACL rupture (Case
j' Si~dy Fig 7-1-1)
-,. The rupture of the primary stabilizer of the knee joint (ACl)
leads to a progressive structural alteration of the knee. A primary objective of the treatment is the prevention of fe-injury of
the knee in the hope of preventing additional ligamentous
Case Study Figure 7-1-1.
of these structures arc malfunctioning or disturbed,
knee joint instability will occur. The ligaments arc
t.he primm)' stabilizer for anterior ancl posterior
translation, varus and valgus angulation, and internal and external rotation of the knee joint (Case
Study 7-1),
F,; et aL (1994) sUlllmarized the functions of the
knee ligaments, The ACL is the predolllinant restraint to anterior tibial displacement. The ligament
accepts 75% or the anterior force at full extension
and an additional 10% (up to 90") of knee nexion,
The posterior crudate ligament 1S the primary restraint to posterior tibial translation: il slIstains 85
to 100% of the posterior force at both 30 and 90" of
knee flexion. The latera) collateral ligament is the
primary restraint to varus angulation and it resists
injuries. meniscal injuries. and possible cartilage degenera·
lion. In this case. the patient firs! completed
a course of
conservative treatment with physical therapy. After 6
months. the subjective inS(clbilily was present during SPOrts
and daily activities such as gail and stair climbing. To com·
pensate for the ACL deficiency, the patient altered his gait
patterns, presenting quadriceps avoidance gait to prevent
the anterior translation of tl1e tibia when the quadriceps
contracts at the midstance phase of the gait (Andriacchi &
Birac, 1993: Berchuck et aI., 1990).
The patient went for surgical treatment. The MRI below
(Case Study Fig. 7-1-2) shows the ACL status after patella
bone-tendon-bone autograft was performed 10 months
post-trauma.
Case Study Figure
7~1·2.
approximately 55% of the applied load at full ext
sion. The role of the lateral collateral ligament
creases with joint flexion as the posterior structu
become lax. The 1l1edial collateral ligament (supe
cial portion) is the primary restraint to valgus (
duction) angulation and resists 500/0 or the exter
valgus load. The capsule. the anterior and poster
cruciatc ligaments, share the remaining valgus lo
Internal rotational laxity seen in the 20 to 40° ran
of knee nexion is restrained bv the medial collate
ligament and the ACL Finally, external rotation l
ity seen in the 30 to 40° range of knee flexion is
strained by the posterior cruciate ligaJllcnt at 90°
knee Ilexion,
In vivo measurement or the normal ACL has b
performed by Beynnon et aL (1992), They place
strain transducer arthroscopic'll)y in the ACL. The
results showed that strain in the ACL was related to
knee Oexion (with the most strain occurring ncar
full extension) and increased with quadriceps contraction. Less strain occurred with co~conlraclions
of both the quadriceps and the hamstring muscle
groups and at greater degrees of knee flexion. This
indicates that muscle contraction and co-contraction
contribute to the stabilitv of the knee joint bv incrcasing thc stiffncss of the joint. Kwak et al. (2000)
studied in vitro the effect of hamstrings and iliotibial band forces on the kinematics of the knee. At various knee Oexion angles, human knee specimens
were tested with different muscle-loading patterns.
The quadriceps muscle force was always present,
and the test was performed with and withollt han)string muscle force and with and without iliotibial
band force. \AJith loading of simultaneous quach·i.
eeps and hamstring muscle force. the tibia translatcd posteriorly and rotatcd externally. The effecI
""" similar for thc iliotibial band simulated forces
but thc cffcct was smallcr.
Many in vitro studies suggest that the hamstrings are important anterior and rotational stab.ilizers of the tibia. In vivo studies have shown that
co-contractions of the quadriceps and hamstring
muscles arc highly present in normal knee joints and
dail.'", activities (Baratta et a1., 1988; Solomonow &
D'Ambrosia, 1994). Thc co-contraction mcchanisms
also increase the knee joint stability in vivo (Aagaard
ct al .. 2000; Markholf et al .. 1978; Solomonow &
D'Ambrosia, 1994). Howevel~ the complex mechanism in vivo of muscle activity as a knee stabilizer,
the extent of protection, and the biomechanical
and clinical imparlance needs futther research
(Grabiner & Wcikcl', 1993).
tendon and it contributes the least to the leng
the quadriceps muscle force lever arm (app
mately 10% of the total length). As the knee i
tended. the patella rises from the intercon
groove, producing significant anterior displace
of thc tcndon. Thc length of thc quadriccps
lever arm rapidly increases with extension up to
at which point the patella lengthens lhe lever
by approximately 300/0.
With knee extension bc"ond 45°, thc Icngth o
lever arm is diminished slightly. \,Vith this dec
in its lever arlll, the quadric.:cps muscle force
increase for the torquc about the knee to rcrnai
same. In an in vitro study of normal knees. Licb
Perry (1968) showcd that thc quadriccps m
force required to extend the knee (he last 15
creased by approximately 60% (Fig. 7-19).
If thc patella is removed from a knec, the I)a
tendon lies closer to the center of motion o
tibiofemoral joint (Fig. 7·20). Acting with a sh
lever arm, the quadriceps llluscle must pro
even more force than is normnlly required for a
tain torque about the knee to be maintained du
75
FUNCTION OF THE PATELLA
The patella serves two important biomechanical
functions in the knee. First, it aids knee extension
by producing anterior displacement of the quach-ieel's tcndon throughollt the entire range of motion.
thereby lengthening the lever arm of the quadriceps
muselc force. Second, it allows a wider distribution
of compressive stress on the femlll- by incrensing the
area of contact between the patellar tendon and the
fcmur. The contribution of thc patella to the Icngth
of the quadl"iceps muscle force lever arm varies
from fullllcxion 10 full extcnsion of thc knee (Lindahl
& Movin, 1967; Smidt, 1973). At full Ilcxion, whcn
the patella is in the intercondylar groove, it pro~
duces littlc anterior displacemcnt of the quadriccps
""','
o
90-60
60-30
30-15
Knee Motion (degrees)
Flexion - - - - EXlension
~~------
Quadriceps muscle force required during knee motio
from 90" of flexion to full extension. Adapted from L
f1. & Perry, 1. (J 968). Quadriceps funccion. An dflacomica
mechanical study using dmplJtMed limbs. J Bone Joint Su
50A, 1535.
,
,.
~~=====
-------------
"-
\\
\
---------
~-
Instant center
Normal
-L=-After Patellectomy
Quadriceps muscle lever arm (represented by the broken
line) in a normal knee and in a knee in which the patella
has been removed. The lever arm is the perpendicular distance between the force exerted by the quadriceps muscle
through the patellar tendon and the instant center of the
tibiofemoral joint for the last two degrees of extension.
The patellar tendon lies closer to the instant center in the
knee without a patella. Adapted from Kaufer. H. (1911). Mechanical function of the parella. J Bone Join! Surg. 53A. 1551.
knee is almost directly above the center of rotatio
of tho patellofemoral joint. As knee Oexion i
creases, the center of gravity shifts further awa
from the center of rotation. thcrcb.y greatly increa
ing the Ilexion moments to be counterbalanced b
the quadriceps musclc force. As the quadricep
muscle force rises. so docs the patellofemoral joi
reaction force (Hungerford & Barry, 1979; Reilly
J\>larlons, 1972).
Knee flexion also influences the patellofemor
joint reaction force by affecting the angle betwee
the patellar tendon force and the quadriceps te
don force. The angle of these two force comp
nents becolTlCs more acute with knee flexion, i
creasing the magnitude of the patellofemoral joi
reaction force (Calculation Box 7-3). Reilly an
Martens (1972) dctcrmincd the magnitude of th
patellofcmoral joint reaction force during sever
dynarnic activities involving val·ying amounts
knee flexion. During IC\'e1 walking, which requir
relatively little knee flexion, the reaction force w
low. The pcak valuc. ill the middle of the stan
phasc when flexion was greatest, was one-ha
body weight.
The joint reaction forcc was much greater durin
activities that require greater flexion. During kn
bends to 90'. this force reached 2.5 to 3 times bod
wcight with the knee Oexed 90' (Fig. 7-21). Throug
3000
the last 45° of eXlen~ion. Full active extension of
such a knee may require as much as 30% more
quach'iceps force than is normally required (Kaufer.
1971). This increase in force may be beyond the capacity of the quadriceps muscle in some patients,
particularly those who have intra~articular disease
or are advanced in age.
1
...... Palellofemoral joinl
reaction force
--- Quadriceps muscle force
:s
2000J
"
!!
0
"1000
STATICS AND DYNAMICS OF THE
PATElLOFEMORAl JOINT
During dynamic activities, the magnitude of the muscle forces acting on a joint directly affects the magnitude of the joint reaction force. In general, the
greater the muscle forces, the greater the joint reaclion force.
In the patellofemoral joint. the quadriceps muscle force increases with knee Oexion. During re~
(axed upright standing, minimal quadriceps muscle
forces are required to counterbalance the small
nexion l1lornents about the patcllofemoral joint because the center or gravity of the body above the
o
20
40
60
80
Knee Flexion (degrees)
Patellofemoral joint reaction force and quadriceps muscl
force during knee bend to 90° (three subjects). Adapted
from Reilly, D. T. g. Martens, M. (1972). Experim.ental analysis o
the quadriceps muscle force and patello/emaral joint reaction
force for various activities. Acta Orthop Scand. 43. 126
Joint Reaction Forces at the Knee in Flexion
Knee flexion influences the patellofemoral joint reaction force by changing the angle between
the patellar tendon and the quadriceps tendon (Calculation Box Fig. 7-3-1, A & B).
The angle between the patellar tendon (P) and the quadriceps tendon (Q) is 35° with the
knee flexed 5° (left top) and 80° with the knee flexed 90° (left bottom). Values for the tendon angles are from Matthews and Associates (1977), who determined the angle roentgenographically after placing two metal wires along each of these tendons.
The pateJlofemoral joint reaelion force with the knee in 5 and 90° of flexion is oblained by
constructing a parallelogram of forces for each situation and using trigonometric calculations.
The pateJlofemoral joint reaction force (J) is the resultant of the hvo equal force components
through the patellar tendon (P) and the quadriceps tendon (Q). As the angle between these
force components becomes more acute with greater knee flexion, the resultant joint reaction
force (J) becomes larger. Adapted lrom Wikrorin Cv.H. & Nordin, M. (1986). Introduction to
Problem Solving in Biomechanics (pp. 87-129). Philadelphia: Lea & Febiger.
o
,,/
p
,,'"
//
P------------/
1000 N
J ~ 601.409
J2 '" Q2 .... p2 _~ 2PQ. cos 35~
Tibio femoral flexion 5"
J = )361695
~ 601.41
J
o
I
1000 N
!---J--""-'r:
80"
I
I
I
I
I
I
p
I
I
I
10;0 N ---------j·::-;285_61
J2
I Tibio femoral flexion 90"
Calculation Box Figure 7·3·1A.
-------_._----
~~
Q2 -;.. p2
J
~ 11652703
J
~
.~
2PQ • cos
80~
1285_57
Calculation Box Figure 7·3-1B.
--------_..._------_._--- --_._.--- -._------
'"
Extensor Mechanism Injury
30-year-old basketball player had a forceful knee flexion while coming down from a jump. A strong eccen-
of the quadriceps produced abnormally high
" ,0n",lp loads in the patella, leading to a fracture at the infeIn this case, the patella fracture occurred because
forces of the quadriceps overcame the osseous
of the patella. The weakest link was the patella.
picture shows a fracture at the patella accompaf)ied by a significant fracture separation that resulted from
the quadriceps traction force.
Because of the fracture. the extensor mechanism is un=;'.
able to function and extend the knee. It will directly affect
the stability of the pate!lofemoral joint and the distribu-
Co;
tion of the compressive stresses on the femur. At the
same time, the impaired function of the quadriceps decreases the dynamic stability at the knee joint
(patellofemoral and tibioiemoral joints) that is necessary
:J.
for daily activities such as gait and stair climbing.
contact surface area increases in size somew
(Goodfellow et aI., 1976). To some extent, this
crease in the contact surface with knee nexion co
pensates for the larger patellofemoral joint react
force, 11' a tight iliotibial band is present,
patcllofemoral joint force may shih laterally, ca
ing abnormal patellar kinen1atics and load-bear
(Kwak et aI., 2000),
The quadriceps muscle force and the tor
around the patellofemoral joint can be extrem
high under certain circumstances, particula
when the knee is flexed-for instance. whe
basketball plaver suffers a patella fracture a
result of indirect forces played by an eccen
contraction of the quadriceps (Case Study 7
Another extreme situation was observed durin
stud~; of the external torque on the knee p
duced by weight lifting: one subject ruptured
patellar tendon when he lifted a barbell weigh
175 kg (Zernicke et aI., 1977). At the instant
tendon rupture, the knee was flexed 90°;
torque on the knee joint was 550 Nm and
quadriceps muscle force was appro:dt11a
10.330 N.
Because of the high magnitude of quadric
muscle force and joint reaction force during
tivities requiring a large amount of knee flex
patients with patcllofemoral joint deran
ments experience increased pain when perfo
ing these activities. An effective mechanism
reducing these forces is to limit the amoun
knee flexion.
Summarv
1 The knee is a two-joint structure that is co
posed of the tibiofel11oral joint and the pate
femoral joint.
out knee bend, the patellofemoral joint reaction
force remained higher than the quadriceps Illuscle
force. During stair climbing and descent, at the point
when knee flexion reached a maximum of approximately 600, the peak value equaled 3.3 times body
weight.
When the knee is extended, the lower part of the
patella rests against the femur. As the knee is flexed
to 90°, the contact surrace between the patella and
femur shifts cranially in vivo and under weightbearing conditions (Komistek et aI., 2000). The
?:>: In the tibiofemoral joint, surface motion
curs in three planes simultaneously but is grea
by far in the sagillal plane. In the patellofem
joint, surface Illotion occurs simultaneously in
planes, the frontal and the ll-ansverse, and is gre
in the frontal plane.
3 The surface joint motion can be described w
the use of an inslant center technique. vVhen
formed on a non11al knee, the technique reveals
following: the instant center for sllccessive inter
of motion of the tibiofcmoral joint in the sagi
plane follows a semicircular pathway, and the di
tion 01" displacement of the tibiofemoral con
points is tangential to the surface of the tibia, indi
eating gliding throughollt the range o[ motion.
w
The scrc\v-homc mechanism of the tibiofemoral joint adds stability to the joint in full extension. Additional passive stability' to the knee is given
by the ligamentous structure and menisci and the
dynamic stability by' the muscles around the knee.
S The tibiofemoral and patellofemoral joints are
subjected to great forces. Muscle forces have the
greatest influence on the magnitude of the joint rc
action force, which can reach several times body
weight in both joints. Tn the patcllofcmoral joint,
knee flexion also affects the joint reaction force,
with greater knee flexion resulting in a higher joint
reaction force.
w
Although the tibial plateaus arc the main loadbearing structures in the knee, the cartilage.
menisci, and ligaments also bear loads. The menisci
aid in distribt~ting the stresses and reducing the
load imposed on the tibial plateaus.
The patella aids knee extension by lengthening
the lever arm of the quadriceps muscle force
throughout the entire range of motion and allows
a wider distribution of compressive stress on the
femur.
REFERENCES
Aagaard. P., Simons~n, E.B., And~rsen, J.L., et a1. (2000). Antagonist mtlscl~ co-'lctiv<ltion during isokinetic extension.
Scali J J1ed Sci Sports, 10(2),58-67.
Andriacchi,1. P. & Birac, D. (1993). Functionalligamellt tcstin" in the anterior cruciat~ li!.!alll~nt deficient knee. C!ill
Or~JlOp Rei Res, 288 (Alarch), 4~0-47.
Andriacchi, T.P., Kramer. G. ;'vl., & Landon, G. C. (1979).
Three-dimensional coordinate data processing in human
motion analysis. J Biolllcch Eng, /01, 279-283.
Andriacchi, T.P. & Strickland, A.8. (1985). Gait analysis as a
tool to assess joint kinetics. In N. Berme, A.E. Engin.
D.A. COl'reis, et al. (Eds.). Biollli'clulllics o{ Norll/a! and
Pathological J1l111WII Articulating .loints (NATO ASI series,
Vol 93, pp. 83-102). Drodrecht. Netherlands: :',·lartinus
Nijhoff.
Baratta, R., Sololllonow, ivi., ZhOll, 8.1-1., Letson, D.,
Chuinard, R., D'Ambrosia, R. (1988). Muscular co-activation: The role of the antagonist musculature in maintaining knee stability. Alii J Sports Alcd, 16, 113-122.
Berchuck i\'1., Andriacchi, TP., Bach, B.R.. Reider, B. (1990)
Gait adapt ions by patients who h,\\'e a deficient anterior
cruciate ligament. J BOlli' Joillt Surg. i2A, 871-877.
Beynnon, 8., Howe, J.G., & Pope, rVI.H. (1992). The measurements of anterior cruciate ligament strain in vivo. lilt 01'lhop, 16. 1-12.
Dra!!,\nich. L.D .. Andriacchi, T.P.. & Andersson, G.B.J. (19
I~nlcraction betwccn intrinsic knec mechanism and
knee extensor mech'lllislll. ) OnhoJ} Res. 5, 539-547.
Drillis, R., Conlini. R., & Bluestein, \1. (1964). Bod~' segl
par'lmc[(·rs. A stln'ey of 1l1l..'<lStll"elllent tec!miqut.'s.
Lil/lbs, 8. 44.
Frankel. V.H. & Burstl'in, A.I-I. (19701. OnJlOptldic Bi
cfUlllics. Philadelphia: 1.t.'a &- Fcbigcr.
Fr'Hlkel, V.H., Burstein, A.H., l\c Brooks, D.B. (1971). Bi
chanics of internal dl'rangeTw:nt of Lht.' knee. Patho
chanics as dl'tcrmincd b.v an~d~'sis of the instant ccnte
motion. J BOlle Joillt SlIrg. 53..\, 945.
Fu, ElL Harner, CD., Johnsoll, D,I.., ct <11. (1994). Bi
chanics of the kncl' ligamenLs: Basic concl'pts and cli
application. II/Sfr COllrsi' Lt.'ell/re, -t3, 137-148.
Fukuda, 'I{., Takai. 5., Yoshino, N., et al. (2000). Impact
Lransmission of the kncc joint~infltlence of leg align
and the role 01" IllL'niscus and articular carlilage. Cli
Bioll/L'ch, 15, SI6-S2l.
Gardner. T.R .. Ateshian, G.A .. Grclsarner, R.P., et al. (199
6 DOF knec testing de\'ice to d~termine patellar trac
and patellofellloral joint contact area \'ia stereo
lOQramrnl'trv. :Idl' Bio('II'.,; :IS.\1E BED. 28, 279-280
Gr)(ldl"dlow, .I.', llungcrl"o'rd, D.S., & Zindd, \1. (19
Patl'llofenloral joillt llH.'chanics and paLhology. I. F
tional ,uwtorn.v of the patellofL'lllO!';t1 joint. J Bonc
Surg. 5813, 287.
Grabine!', .\-I.D. & Weiker, G.G. (1993). :\ntl.'rior cruciate
ment injury and hamstring CO-<lcli\'ation. elillical
IIlt'ch, 8. 216-119.
Grood, E.S. l\: Suntav. \V.J. (1983). A coordinate svstell
clinical descriptio;l of three dimensional motion~ App
tion to the knee. J Biollll!ch Ellg. 105, 136-144.
He!llIL', 1-1.1. (1990). Biomechanics of lhc patellol"emor.d
and its clinical relevance. Cfill Orthop, 258, 73-85.
Hell"eL, A..J. (1974). AnatolllY and mechanics of mO\'ellle
the knec joint. In A. He]fct (Ed.), Disorders o(the Knc('
1-171. Philadelphia: J.B. Lippincott.
Holden, J.P., Chou. G.. &- Stanhope, S.J. (1997). Change
knee joinL function o\"er a \~'ide range of walking spc
Cfillica! Bio/llL'ch,12(6L 375-382.
Hungerford, D.S. &- BarT\', ;\:1. (1979). Biomechanics of
patdlofelTloral joint. Cli,l OnhoJ!, Ic+c+, 9-1 S.
Kapandji, l.A. (1970). Thc knl'c. In LA. Kapandjii (Ed.l.
Physiology oFthe Joillts (Vol 2, pp. 72-135). Paris: Edi
\1<1 loi ne.
Kaul"er, H. (1971). \Iechanical function of tllt' patella . .l
Joill! Surg. 53A. ISS I.
Kellis, E. & BHltzopCllJlUS, V. (1999). In \'i\'o delerminatio
the patella and Iwmstrings rnonh.'nt arms in adult m
using videofilloroscopy during submHxinwl knee l':\tl'n
and flexion. Cfinical Biolliech. I -t, 118-124.
Kcttelkamp, D.B" Johnson, R.1., Smidt, G.L .. Chao,
Walker, rv1. (1970). An c1ectrogoniomctric stud.\" of
motion in normal gait. J BOlli': Joilll Sur.>;, 52:\, liS.
Kettelkamp, D.B. &- Jacobs, A.\V. (1972). Tibiofclllora! con
area-dL'termination and implications. J BOIIL' Joil/t S
5-tA. 349.
Komistek, R.D., Dennis, D.A., ;\labe, J.A., d HI. UOOO). A
vh'o determination of pHtellol"ernora! c(ll1tact posit
Cfillical Biol/lcch, 15,29-36.
Kroelller. K.I-I .. M;llTas. \V.S ...\1cGlothin. J.D .. 1..'1 al. (1990).
On the IlH.'ilsun:nh.'nts of human stn:ngth. 11111 Imll/strial
rr"OIUl111ic!I·. 6. 199-210.
Kwak~ 5.0.. ,\hmad, C.S .. Gardner. T.R .. 1..'1 al. (2000). Ham·
Sirings and iliotihial forces affect knee ligamellts and contact palh:rn.l O,.,ltop ReS. IS, 101-IOS.
LaTllorc;IU.\:. L.. (1971). Kinematic IllC:lsurt:TllcnIS in Ihl.· stlldy
of hurn:tll w;:lIking. Biomcch;lnics Lah. Uni\'t.'rsiIY of (;tliforni:l, S'lll rr~ncisco. B//II Prostlteth' Res, Sp1971, 10-15.
L:ltlbenth;d. !\..N .. Smidt, G.L., & !\.ctLdkamp, D.B. (1971)."\
quantitative analysis of knee motion during activities of
dail\' living. I'II\'S Til!!,., 52. 3-L
LC\'I.·ns: '\,S"Yllll1l~lI1. V.T., & Blosser. J.'\. (1948). Trans\'crsc
rot;ltion of the seglll"fHs of the lower cXlr('mity in locomotion, 1 BOlli! 10iut Sur..!.:. 30A. S59.
Licb. FJ. ~ I)erry. J. (I968). Quadriceps function. An ;UHllOI1lical ;lnd mcch:lllical sludy using ;lIl1putaled limbs. 1 Bmw
loil/t Surg, 50~\, 1535.
LiIH.bhl, 0 . .& Mm·in,.-\. (1967). TIJ(,' 1lH.'chanil':-; of l'xt...'nsion
of the kncc joint. .·\eta 01'1/1011 Scaml, 38, 226.
~;:;II'kholr. K., Gr:lff-R;l<lford, f\., &. Amstul'I.. H. (19iS). In \'i\'O
knce slllbility. 1 BOlli' 10ill/ SlIl'g. 60.'1. 664-674.
~laISllmot(), 1-1.. Si..'i..'dholll, B.B.. Suda. Y., C"t <:II. (2000). Axis of
tibial rot.Hioll and its change with flcxion angle, C/ill Or·
.
'),Op, 371. 178-182.
',\l<1ltl1(:\\,s, L.S .. Sonstc:gard. D.A .. &. I-lenke, .I.A. (1977). Lond
h"arillg characterislics of lhe palellofemor;d joint. ,·\era
Onhop S,·a//(I. 48.311.
M'HTis('n.J.B. (1970). The mechanics of the kIWi..' joint in relation [0 nonn:'l! walking. J Bicol/cch. 3. 51.
!'\i1.P., Droughl, :\.B .. !\.ory. R.C. (1964). \Valking 1':11ll'::l'nS of l1orl11:d men. 1 130111..' Joillt S!l!·.~, -I6A, 335.
N. & Nordin. rio'\. (1999). FlIlId(/I!U'IIUJ!S oj' 13iOlllt'Eqllilihl'illlll, JIOlio/l, (/lId lJeji>rlllatioll (2nd cd.).
Ncw York: Springcr-V"r1ag.
~crr.v, 1.. Norwood, L., & House. K. (1977). Klll'C po.o;tUl'\.' ,llld
biceps and semimembranosus muscle action in fUllning
<'Ind cutting (an EMG study). hailS Ortlwl) UL'S S(lC 2, 258.
Ramsi..'y, O.K. & Wn.'tl'llhcrg, P.E (1999). Biolllcchallit,:-; of tht:'
kn~c: Melhodological consid('r;ltions in Ihl' in \'i\'o kinl'matic an:dysis (~f II\(' tibiofl'moral and patellofclIlor;d
joint. RC\'i(:w pnp('r. Clillical Biomecll. 1-1, 595-611.
Reilly, 0.1'. &. "'Iancus. ,\'1. (19i2). E.\:pel'imclll;.'I1 ;Inalysis
thl' qlwdriceps musclc force and patcllof"moral joint rea
tion force for variolls 'lctivitic-s. :\ew Orrhop Scalld. 4
126.
Reulc;1ux. r. (1876). In The KillCl1lalio uf Jlachillc:ry: Outli
of II Theory of .'-1achillcs. London: ~l[lcmi lIan.
Sccdhom. B.s.. Do\\'son, D., &. Wright. V. (19i4). The 10;
bl:'aring function of Ihl' menisci: r\ prdiminary study.
0.5. lngwl'rsl'n, ct :II. (Eds.), Thc h//CC loill/: f?en.'1/t .'
\'(l/ICC.~· ill Basic Research (wd Clinica! .'lsp(~rlS (I'p. 37-4
r\mS!l:rdam: Exccrpta Medica.
Seh'ik. G. (1978). Roentgcn slC'l'cophotogr.lllllllclry in Lun
Sweden. In .-\.M. Coble!l'l. & R.E. j·Jl'lTon (Eds.), Applic
/jolls
flUlIltll1 Biosten:mJII..' lru:s. fIroe 81'/£ (166) (p
184-189).
Sdvik, G. (1983). Roentgen stcrl.'ophotogramllleiry in 0
thop;:ledics. In R.E. HelTon (Ed.), BioslereOluetric:s 'S
Pro" SP/E (36/) (pp. li8-18S).
Smidt. G.L. (1973). Biomechanical :-nnlysis of knee nexi
;lUd c.\:wnsioll_ 1 Bioll/celt, 6. 79.
Solomonow, M. &. D'r\mbrosia. R. (1994). Nellr:!1 l'1'f1ex ;I
;tnd muscle control of knl'c sl;\bilil~' :lIld motion. In W
Scott (Ed.), The Knce (pp. 107-120). New York: Moshy.
Vl'di, V., Williams, A" Tennant, S.L ('I :II. (1999). Mcnisc
rnon::mcnt. An ill-vivo study llsing dynamic i\-lRL J BO
loim SlIrf:!.. 8IB(J), 37-41.
Winll'l', D.A. (1990). In Bio!llee/ullIics flud Motor CO/ltrol
1'11111/1111 Be!wviour (2nd ed.). Ncw York: John Will'Y
or
SOilS.
Wiktorin. C.v.H. &. Nordin, ~\l. (1986). 1,II/,o(//I(:t;OI/ [() Prohll.
So/l'illg ill Biollll..'challics (Pl'. 87-129). PhihlClclphia: Lea
Fchig('l'.
Wilson. S.A., Vigorila, V.J., &. SCf.HL. WN. (1994). Anatomy.
N. SCOIt (Ed.), The Kllee (po 17). Philadelphia: ~'Iosby.
Wrclenberg, P., Nemeth. G., Lamontagne. M.. ct <II. (199
P;lssivc knee muscle momcnt <trillS meas.ured in vivo w
MRI. Clilliea! Biol1lcclt, I 1(8}, 439-446.
Yu, B., Stuart, 1\-1.1 .• Kicnb;ll'chcr, T., et .d. (1997). Valg
\·'II·US motion of the knee ill norm.1I lc"d w;dking [lnd st
climbing. Clillical Biolllcch, 12(5), 2S6-293.
Zcrnickc, R.E, G;lrhammer. J., & Jobe. FW. (1977). Hum
pntdlar Icndon rupture. 1 801lt' loillt S1Irg, 59;\, 179-18
Biomechanics
of the Hip
Margareta Nordin, Victor H. Frank
Introduction
Anatomical Considerations
The Acetabulum
The Femoral Head
The Femoral Neck
, /f
Kinematics
Range of Motion
Surface Joint Motion
Kinetics
Statics
Dynamics
Effect of External Support on Hip JOll1l Reaction Force
Summary
References
Ie
~,
Introduction
The hip joint is one of the largest and most SLable
joints in the body. In contrast to the knee, the hip
joint has intrinsic stability provided by its relatively
rigid bali-unci-socket configuration. (t also has a
mobility, which allows normallocomogreat deal
lion in the performance of daily activities. Derangements of the hip Can produce altered stress distributions in the joint cartilage and bone, leading to
degenerative arlhritis. Such damage is further po~
tcntiated by the large forces borne by the joint.
or
-~
Anatomical Considerations
The hip joint is composed of the head of the femur
and the acetabulum of the pelvis (Fig. 8-1). This articulation has a loose joint capsule and is surrounded by large. strong muscles. The construction
of this stable joint allows for the wide range of 1110tion required for normal daily activities such as
walking, sitting, and squatting. Such a joint must be
precisely' aligned and controlled,
THE ACETABULUM
The acetabulum is the concave component of the
ball-and-socket configuration of the hip joint. The
acetabular surface is covered with articular cartilage that thickens peripherally (Kempson et aI.,
1971) and predominantly laterally (Rushfeld et aI.,
1979). The cavity of the acetabulum faces obliquely
forward, olitwarcL and downward. The osseous acetabululll in the hip is deep and provides substantial
static stability to the hip. A plane through the circumference of the acetabulum at its opening would
intersect with the sagittal plane at an angle of 40~
opening posteriorly and with the transverse plane at
an angle of 60~ opening laterally. The acetabular
cavity is deepened by the labrum, a flat rim of fibro~
cartilage, and the transverse acetabular ligament
(Fig. 8-2). The labrum contains free nenle endings
and sensory end organs in its superficial layer,
which may participate in nociceptive and proprioceptive mechanisms (Kim & Azuma, 1995).
The unloaded acetabulum has a smaller diameter
than the femoral head (Greenwald & Haynes, 1972)
(Fig. 8-3). The acetabulum deforms about the
femoral head when the hip joint is loaded. It undergoes elastic deformation to become congruous
with the femoral head, and contact is made abollt
the periphery of the anterior, sllperi()l~ and posterior
The hip joint (front view) 1. External iliac artery. 2. Ps
major muscle. 3. Iliacus muscle. 4. Iliac crest. S. Gluteu
medius muscle. 6. Gluteus minimus muscle. 7. Greate
trochanter. 8. Vastus lateralis muscle. 9. Shaft of femu
10. Vastus medialis muscle. 11. Profunda femoris vess
12. Adductor longus muscle. 13. Pectineus muscle. 14. M
circumflex femoral vessels. 15. Capsule of the hip join
16. Neck of femur. 17. Zona orbicularis of capsule.
18. Head of femur. 19. Acetabular labrum. 20. Rim of
etabulum. Reprinted with permission (rom McMinn, R.H.
Huchings. R.H.R. (988). Color Atlas of Human Anatomy
ed., p. 302). Chicago: Year Book Medical Publishers, Inc.
arLicular surface of the acetabulum (Konrath et al..
1998). Load distribution of the acetabulum was
studied in vitro in human specimens (Greenwald &
I-(aynes, 1972; Konrath et 'II., 1998). Joint reaction
forces were simulated to physiological levels. The
loading pattern of the acetabulum is shown in Figure 8-3. Removal of the transverse acetabular ligament and labrum sequentially did not affect the
loading pattern of the acetabulum significantly
(Konrath et 'II., 1998).
Superior
Anterior
Intact
THE FEMORAL HEAD
The femoral head is the convex component of the
ball-and-socket configuration of the hip joint and
forms two-thirds of a sphere. The articular cartilage covering the femoral head is thickest on the
medi,al-central surface and thinnest toward the periphery. The variations in the cartilage thickness
result in a different strength and stiffness in various regions of the femoral head (Kempson et 'II.,
loading pattern of a human acetabulum in vitro with
tact labrum and transverse acetabular ligament. Note:
pattern was grossly unchanged after removal of the tr
verse acetabular ligament. or the labrum, or both, and
therefore those patterns are not displayed. Adapled fro
Konrath. GA.. Hamel. A.I.. Olson. S.A.. et elf. {l998). The ro
the acetabular labrum and the trdnSverse clCeiabtJ/dr ligame
load transmission of tilt, hlp. J Bone Joint Surg. 80/'.,(12),
.--I
Transverse
acetabular
ligament
1781-1788.
J 971 l. Rydell (1965) suggested I hat most load
transmitted in the.: femoral head through the su
rior quad,-ant. VOll Eisenharl-Rothe et al. (19
demonstrated in an in vitro study that the ma
tude or load influenced tlte loading patlerll on
femoral head. At smaller load:s, the load-bea
area was concentrated at the periphery of the
nate surface of the femoral head. and at hig
loads to the center of the lunate and the ante
and posterior horns. It is still not known exa
how the stresses in vivo on the normal fem
head arc distributed. bUl indications from in
measurements with an instnll11Cnted prosth
head show that the anterior and medial lunat
transmitting most 01" the load during daily ac
ties (Bergmann et al .. 1993. 1995).
THE FEMORAL NECK
Schematic drawing showing the lateral view of the acetabulum with the labrum and the transverse acetabular ligament intact. Adapred (rom KOflrarh, G.A., Helmel. Ai.. Olson,
s.A., et al. (998). The role of (he acetabular labrum and tile
transverse acetabular ligamenr in (oad transmission of the hip.
J Bone Joint Surg. 80A(12), 1781-1788.
The fen10ral neck has two angular relations
with the femoral shaft that arc important to
joint function: the angk of inclinalion of the nec
the shaft in the frontal plane (the neck-to-shaft
gie) and the angle of inclination in the transv
'plane (the angle of anteversion). Freedom of rno
of the hip joint is facilitated bv the neck-to-shaft
gle, which offsels the femoral shaft from the p
laterally. in 11'10$1 adults. this angle is approxima
125~, but it can vary' from 90 to 135 . An angle exceeding 125' produces a condition known as coxa
valga; an angle less than 125'" results in coxa vara
(Fig. 8-4). Deviation of the femoral shaft in either
\vay alters the Force relationships about the hip joint
and has a nontrivial effect on the lever arms to muscle force and line of gravity'.
The angle of anteversion is formed as a projec~
l.ion of the long axis of the femoral head and the
transverse axis of the femoral condyles (Fig. 8~5).
In adults, this angle averages approximately 12"',
but it varies a great deal. Anteversion of more than
12" causes a portion of the femoral head to be uncovered and creates a tendency tc}\vard internal rotation of the leg during gait to keep the femoral
head in the acetabular cavity. An angle of less than
12" (rclroversion) produces a tendency toward ex~
ternal rotation of the leg during gail. Both anteversion and retroversion arc fairly common in children but arc usually outgrown.
The interior of the femoral neck is composed of
cancellous bone \vith trabeculae organized into
medial and lateral trabecular systems (Fig. 8-1, 86), The fact that the joint reaction force on the
femoral head parallels the trabeculae of the medial
Neck-to-Shaft Angle
i
Coxa vara
angle <.: 125"
1DIIlI
Normal
125C,
angle
Coxa valga
angle> 125'
_
The normal neck-to-shaft angle (angle of inclination of
the femoral neck to the shaft in the frontal plane) is approximately 125". The condition in which this angle is less
than 125' is called coxa vara. If the angle is greater than
125·, the condition is called coxa valga.
0------------------
Neck
Head
Greater
trochanter
Lateral
condyle
Medial
condyle
Top view of the proximal end of the left femur showing
the angle of anteversion, formed by the intersection of th
long axis of the femoral head and the transverse axis of
the femoral condyles. This angle averages approximately
12' in adults.
system (Frankel, 1960) indicates the importance
the s}'stem for supporting this force. The epiphy
seal plates are at right angles to the trabeculae
the medial system and are thought to be perpen
dicular to the joint reaction force on the femor
head ([nman. 1947). It is likely that the lateral tr
becular system resists the compressive force on th
fcmoral head produced by contraction of the ab
ductor muscles-the glutcus medius, the gluteu
minimus, and the tensor fasciae latac. Thc th
shell of cortical bone around the superior femor
neck progressively thickens in the inferior region
\Vith aging, the femoral neck graduall.v undergoe
degenerative changes: the cortical bonc is thinne
and cancellated and the trabeculae arc gracluall}! r
sorbed (see Fig. 2-50). These changes may predi
pose the femoral neck to fracture. It is noteworth
that the fernoral neck is the most cOllllllon fractu
site in elderly persons (Case Study 8-1).
Kinematics
In considering thc kinematics of the hip joint, it
useful to view the joint as a stable ball-and-sock
configuration wherein the femoral head and aceta
ulum can move in all directions.
three planes. Measurements in the sagittal
during level walking (Murray, 1967) showed
the joint was maxirrwlly tlexed during th
swing phase of gail, as the limb moved forwa
heel strike. The joint extended as the body m
fOl\vard at the beginning of stance phase.
ll)um extension was reached at heel-oil. The
reversed into flexion during swing phase and
reached maximal Ocxion, 35 to 40', prior to
strike. Figure 8-81\ shows the pattern of hip
motion in the sagittal plane during a gail cyc
----------
Femoral Intertrochanteric Fractures
A
.,
n 80-year-old woman falls from her own height af
losing her balance. She presented with sharp pain
her hip and an inability to stand and walk by herself. Sh
is transported to the E.R. and after a careful examinatio
and x-ray evaluation, a right intertrochanteric fracture is
d,iagnosed.
Roentgenogram of a femoral neck showing the medial
and lateral trabecular systems. The thin shell of cortical
bone around the superior femoral neck progressively thick·
ens in the inferior region .
•
RANGE OF MOTION
Hip motion takes place in all three planes: sagittal
(flexion-extension), frontal (abduction-adduction),
and transverse (internal-external rotation) (Fig.
8·7). Motion is greatest in thc sagittal plane, where
the range of ncxion is from 0 to approximatcly 140"
and the range of extcnsion is from 0 to 15~. The
range or abduction is fTom 0 to 30", whereas that of
adduction is somewhat less, from 0 to 25'. External
rotation ranges from 0 to 9W degrees and internal rotation from 0 to 70' when the hip joint is flexed. Less
rotation occurs when the hip joint is extended because of the restricting function of the soft tissues.
The range of motion of the hip joint during gail
has been measured e1ectrogoniometrically in all
;'
_~o_,
~:,,;:h-
The radiograph illustrates a right femoral intertrochanteric unstable fracture with separation of the
lesser trochanter. The image shovvs osteoporOlie changes
characteristic of the aging process. The decrease in the
bone mass at the femoral neck leads
(0
reduced bone
strength and stiffness as a result of diminution in the
amount of cancellous bone and thinning of conical bone
increases the likelihood of a fracture at the weakest level
During the fall. the magnitude of the compressive
forces at the femoral neck overcame its stiffness and
strength. In addition. the tensile forces produced by pro
tective contraction of muscles such as the iliopsoas gen
ated a traction fracture at the lesser trochanter level.
:t: ./
i;:
"
Extension
Abduction
B
Adduction
External
rotation
c
D
Internal
rotation
E
Movements of the hip joint. A. Flexion-extension. B, Abduction. C. Adduction. D. External
rotation. E. Internal rotation .
•
't.
'T·- .. ,
allows a comparison of this motion with that of the
knee and ankle.
Motion in the frontal plane (abduction-adduction) and transverse plane (internal-external rotation) during gait (Johnslon & Smidt. 1969) is illustrated in Figure 8-8B. Abduction occurred during
swing phase, reaching its maximum just after toeolT; at heel strike, the hip joint reversed into adduetion, which continued until late stance phase. The
hip joint was externally rotated throughout the
swing phase, rotating internally just before heel
st.-ike, The joint .-emained internally rotated until
late stance phase, when it again rotated externally.
The average ranges of motion recorded for the 33
normal men in this study' \vere 12' for the frontal
plane and 13' For the transverse plane.
As people age, they use a progressively smaller
portion of the range of motion of the lower extrem~
ity joints during ambulation. Murray and co\\,/orkers
(1969) studied the walking pallerns of 67 normal
men of similar weight and height ranging in age
from 20 to 87 years and compared the gait patterns
of o.lder and younger men. The diffcrences in the
sagittal body positions of the two groups at the instant of heel strike are illustrated in Figure 8-9. The
older mcn had shorter strides, a decreased nmge of
hip flexion and extension, decreased plantar flexion
of the ankle. and a decreased heel-to-noOJ" angle of
'-.
,.
the tracking limb; they also showed reduced do
nexion of the ankle and diminished elevation of
toe of the forward limb,
The range of motion in three planes during c
mon daily activities such as tying a shoe, sit
down on a chair, rising [Torn it, picking lip an ob
from the floor, and climbing stairs \vas measu
electrogoniomctrically in 33 normalmcn by John
and Smidt (1970), The mean motion during th
activities is shown in Table 8-1. Maximal Illatio
the sagittal plane (hip flexion) was needed for t
the shoe and bending down to sqUal to pick lip
object rrom the noor. The greatest motion in
frontal and transverse planes was recorded du
squatting and during shoe tying with the foot ac
the opposite thigh, The values obtained for t
common activities indicate that hip flexion o
least 120 abduction and external rotation of atl
20" are necessary for caITying out daily activitie
a normal manner.
G
SURFACE JOINT MOTION
Surface Illotion in the hip joint can be considere
gliding of the femoral head on the a<;etabulum.
pivoting of the ball and socket in three planes aro
the center of rotation in the femoral head (estim
at the center of the femoral head) produces this
--Hip joint
70
----. Knee joint
........... Ankle joint
60~
50
'"
w
w
:s
c
0
1
/
20
~
10
u:
C
,
~
c
10 ...
0
'.'!l"
c
x
LU
"
"
.••....
,
5
•,
1
I
.
..
w
w
,
Adduction
w
:s
c
.
\
5
Q,
\
.., /
.....
(j)
\
I'
........
o i==""1--~
,
'"
Q
Internal
rotation
;;
~
,
"'i...---' I
5
...-
External
rotation
100
Stance phase
60,
Swmg phase
--.l
51:::::=~....1
£Q. 04
......
................
20
30
100
Abduction
\
'I
./
0
'0
\
1
,,
30~
c
,2 'x
w
;;
:
,,•
40~
~
C>
.
I.
,-',
,, ,,
,,
,
100
Stance
phase
60
Swing
phase
100
Percentage of Cycle
Percentage of Cycle
B
A
A. Range of hip joint motion in the sagittal plane for 30 normal men during level walking, one gait cycle. The ranges of
motion for the knee and ankle joints are shown for comparison. Adapted from ivlurray, ~J1.p. (1967). Gait as a tolal pattern of
movement. Am) Phys iVIed, 46, 290. B, A typical pattern for
Differences in the sagittal body positions of older men
(left) and younger men (right) at the instant of heel strike.
The older men showed shorter strides, a decreased range
of hip flexion and extension, decreased plantar flexion of
the ankle, and a decreased heel"to-floor angle of the
tracking limb; they also showed less dorsiflexion of the an"
kle and less elevation of the toe of the forward limb.
Reprinted with permission from Murray, MP., Kory, R,C, &
Clarkson, 8,H. (1969). Walking patterns in flealthy old men.
Gerontal, 24, /69-178.
range of motion in the frontal plane (top) and transver
plane (bottom) during level walking, one gait cycle. Ad
from Johnston, Pl. C. S Smiclt, G.L. (J 969). /vleasuremcn[ of h
joinr morion during waf.l-:.ing_ EvaluMion of an electrogoliiom
method J Bone Joint Sur~J, 51 A 1083
ing of the joint surfaces. If incongruity is prese
the femoral head, gliding may not be parallel or
gential to the joint surface and the articular cart
may' be abnormally compressed or distracted. In
center analysis by means of the Reulcaux me
cannot be performed accurately in the hip join
cause motion takes place in three planes simul
ously. Locating the center of rotation of the hip
is essential for prosthetic surgery' of the hip to re
struct an optimal lever arm of the gluteus me
muscle (Fess)' et aI., 1999),
Kinetics
Kinetic studies have demonstrated that substa
forces act on the hip joint during simple activ
(Hurwitz & Anclriacchi, 1997, 1998), The faclo
volved in producing these forces must be lI
stood if rational rehabilitation programs arc
Mean Values for Maximum Hip Motion in
fhree Planes During Common Activities
Plane
of Motion
Recorded Value
(Degrees)
Tying shoe with
foot on floor
Sagittal
Frontal
Transverse
124
19
1
Tying shoe with
Sagittal
Frontal
Transverse
1
foot across
opposite thigh
Sitting down
on chair and
rising from sitting
Frontal
Transverse
Activity
Stooping to obtain
object from floor
Squatting
Sagittal
Sagittal
Frontal
Transverse
Sagittal
Frontal
Transverse
J'>.scending stairs
Sagittal
Frontal
Transverse
Descending stairs
Sagittal
23
33
104
20
117
21
122
28
26
67
16
18
36
i'.ile,)1l for 33 normal men. Data hom Johnston, R.C. & Smidt. G.L.
i.i970), Hip motion measurements for selected activities of daily living;
(lin Orrhop, 72, 205
developed for patients with pathological conditions
of the hip. The abductor muscle group (the glutcus
medius and minimus muscles) is the main stabilizer
during one-legged stance (Kumagai et aI., 1997;
University of California, 1953)
STATICS
During a two-leg stance, the line of gravity of the
superincumbent bod:v passes posterior to thc pu~
bic s)'mphysis, and, because the hip joint is stable,
an erect stance can be achieved \vithout muscle
contraction through the stabilizing effect of the
joint capsule and capsular ligaments. With no
muscle activity to produce moments around the
hip joint, calculation of the joint reaction force becomes simple: the magnitude of this force on each
femoral head during upright two-legged stance is
one half the weight of the superincumbent body.
Because each lower extremity is one sixth of body
weight, the reaction force on each hip joint will
one half of the remaining two thirds, or one thi
of body weight; however, if the muscles surroun
ing the hip joint contract to prevent swa,ying an
to maintain an upright position of the body (e.
during prolonged standing), this force increases
proportion to the amount of muscle activity.
vVhen a person changes from a two-leg to a singl
leg stance, the line of gravit:v of the superincumbe
body shifts in all three planes, producing momen
around the hip joint that must be counteracted
muscle forces and thus increasing the joint reacti
force. The magnitude of the moments, and hence t
magnitude of the joint rcaction force, depcnds on t
posture of the spine, the position of the non-wcigh
bearing leg and upper extremities, and the inclin
tion of the pelvis (McLeish & Chamley, 1970). Figu
8-10 demonstrates how the line of gravity in t
frontal plane shifts with four different positions
the upper body and inclinations of the pelvis: stan
ing with the pelvis in a ncutral position, standi
with a maximum tilt of the upper bod)' over the su
porting hip joint, sianding with the upper body ti
ing away from the supporting hip joint, and standi
with the pelvis sagging away' from the supporting h
joint (Trendelenburg's test). The shift in the grav
line, and hence in the length of the lever arm of t
gravitational force (the peq)endicular distance b
tween the gravity line and the center of rotation
the fernoral head), influences the magnitude of t
moments about the hip joint and, consequently, t
joint reaction force. The gravitational forcc lever ar
and the joint reaction force are minimized when t
trunk is tilted over the hip joint (Fig. 8-108).
Two methods arc used for deriving the magnitu
of the joint reaction force acting on the femor
head: the simplified free-body technique for cop
nar forces and a mathematical method utilizi
equilibrium equations. The simplified free~bo
technique for coplanar forces was described in d
tail in Chapter 7, in Calculation Box 7.1. This tec
nique is used in the hip to cstimatc the joint rea
tion force in the frontal plane acting on the femor
head during a single-leg stance with the pelvis in
neutral position (Calculation Box 8-1). The seco
method is a mathematical calculation of the jo
reaction force on the femoral head using equil
rium equations for a single-leg stance with t
pelvis level (Calculation Box 8-2).
To understand and solve the equations it is ne
essaI)' to indicate first the location of the extern
forces acting on the body during the single-l
Roentgenograms utilizing a plumb line (black line) show
that the line of gravity shifts in the frontal plane with differ·
ent positions of the upper body and inclinations of the
pelvis. A, The pelvis is in a neutral position. The gravity line
faUs approximately through the pubic symphysis. The lever
arm for the force produced by body weight (the perpendicu·
lar distance between the gravity line and the center of rotation in the femoral head) influences the moment about the
hip joint and hence the joint reaction force. B, The shoulders
are maximally tilted over the supporting hip joint. The grav·
ity line has shifted and is now nearest the supporting hip.
Because this shift minimizes the lever arm, the moment
about the hip joint and the joint reaction force are also minimized. C, The shoulders are maximally tilted away from the
supporting hip joint_ Again the gravity line has shifted toward the supporting hip, thus decreasing the joint reactio
force. D. The pelvis sags away from the supporting hip join
(Trendelenburg's test). The shift in the gravity line is simila
to that in C. (Courtes}' of 10hn C Baker. IvI.D., Case vVesrem R
serve Unive(sir}~ Cleveland, OhiO} Note: In B, the antalgic gait
illustrated, which lowers the load on the head of the femu
but alters the load line to a more vertical position. Followi
arthroplasty for arthritis, the abductor muscles are weak a
atrophic as a result of the disease process and the surgery.
External support such as a cane should be used until the a
ductor muscles are rehabilitated. The best indication for a
rehabilitated abductor muscle is the lack of limping.
on a free·body diagram (Calculation Box Fig.
1)_ Because the body is in equilibrium (i.e" the
of the 11''lOmenls and the sum of the forces both
zero), the ground l'eaction force is equal to the
'Q:l-,witatwllal force of the body, which can be divided
1WO components, the gravitational force of the
leg (equal to one-sixth body weight) and the
force (equal to five-sixths body weight).
body is divided at the hip joint into two
h-"e-DC)(lIeS- The main coplanar forces and moments
on these free-bodies must be determined.
upper free-body is considered first (Calculalion
;1j~IJOX Fig. 8-2-2). In this free-body, t\\'o moments are
required for stability. The moment arising from th
superincumbent body weight (equal to % 'IN) mu
be balanced by a moment arising from the force
the abductor muscles_ The rorce produced by the s
perincumbent hody weight (% W) acts al a diSlan
or b rrom the center of rotation of the hip (Q), thu
producing a moment of % Vt/ times b. The force pr
duced by the principal abductor, the glute
medius, designated as A, acts at a distance of c fro
the center of rotation, producing a counterbalan
ing rnomcnt of A limes c. For the body to remain
moment equilibrium. the sum of the moments mu
equal zero. In this example, the moments actin
-----------------Simplified Free-Body Technique for Coplanar Forces
The stance limb is considered as a free-body, and a free-body
lion, simplifying assumptions are made in determining the di-
diagram is drawn. From all of the forces acting on the free-
rection of this force (Mcleish & Charnley, 1970). Furthermore.
body, the three main coplanar forces are identified as the
forces produced by other muscles aCiive in slabilizing the hip
force of gravity against the foot (the ground reaction force),
joint are not taken into account. The joint reaction force (])
\,.'/hich is transmitted through the tibia to the femoral
has a known point of application on the surface (lunate) of
condyles; the force produced by contraction of the abductor
the femoral head bur an unknown magnitude, sense, and
muscles; and the joint reaction force on the femoral head.
line of application.
The ground reaction force ('IN) has a known magnitude equal
to five-sixths of body weight and a known sense, line of ap-
.'.;
The magnitudes of the abductor muscle force and the
jOint reaction force can be derived by designating all thr~e
plication. and point of appliCc1tion. The abdu([or muscle force
forces on the free-body diagram (Calculation Box Fig. 8~_1·
(A) has a known sense, a known line of application, and
and constructing a triangle of forces (Calculation Box Fig:
point of application estimated from the muscle origin and in-
8-1-2). The muscle force is found to be approximately two
n
sertion on a roentgenogram but an unknown magnitude. Be-
times body weight, whereas the joint reaction force is some-
cause several muscles are involved in the action of hip abduc-
what greater.
Force J
Intersection
point
J'
//
Force A
I
2W/
/
,
1
!
t
Force W
Calculation Box Figure 8-1-1.
,._wt/
I
I
I
Force J
2_75 W
Calculation Box Figure 8·'-2.
------------------------------------ -----
&--------------------------------------~----------
,.- . ,
-:-.-_ ..
External Forces Acting on the Body in
Equilibrium During a Single-Leg Stance
Calculation Box Figure 8-2-1 shows external fOKes acting on
the body in equilibrium during a single-leg Slance. The ground
reaction force is equal to body weight (W). The gravi!ational
force of the stance leg is equal to one-sixth of body weight;
Calculation Box
Figure 8-2·1.
the remaining force is equal to five-sixths of body weight.
w
The internal forces acting on the hip joint are founel by sepa-
rating the joint into an upper and lower free-body; the Lipper
free-body is considered first. In this free-body, two moments
b
5/6 W :
:
-_._"
are required for stability. Moment equilibrium is attained by
the prOduction of two equal moments. A moment arising
Calculation Box
Figure 8-2·2.
from the force of the abductor muscle (A) times abductor
force lever arm (e) counterbalances the moment arising from
the gravitational force of the superincumbent body (5/6 W)
times gravitational force lever arm (b), which tends to tilt the
pelvis away from the supporting lower extremity. Q, center of
rotation of the hip joint.
A = 2W
Ax
=:
A
·sin30~
Ax = O.5A
Ay
:=
~
W
A·cos30'"
A y = 0.8 A = 1.7 W
Ay
30"
Calculation Box
Figure 8-2-3.
Force A is equal to two times body weight and has a direction
of 30' from the vertical. The magnitudes of its horizontal (A,)
and vertical (A) components are found by vector analysis.
Perpendiculars are drawn from the tail of A to a horizomal
and a vertical line representing A. and A" respectively. A, and
A, can then be scaled off. Allernatively. trigonometry is used
to find the magnitudes of the components.
.._--------._,
._--~._._-_._-_
Ax
clockwise arc arbilraril~' considered to be positive
and lhe counterclockwise moments arc considered
be negative. Thus,
(Y.,W x b) - (A x c) = 0
A=
YoW x b
-'----
c
To solve for A it is necessary to find the values of
.~.'
and c. The gravitational force lever ann (b) is
roentgenographically. Because the cenler of
gravity must lie over the base of sUppOrl, a plumb
line intersecting the heel can be extended upward; a
line drawn from the center of rotain the femoral head (Q) to the line represents
"""'\IlOe b. The muscle force lever arm (c) is similarl~! found by identifying the glutcus medius muscle on a roentgenogram (Nemeth & Ohlsen, 1985,
1989) and drawing a perpendicular line from the
center of rotation 01" the femoral head to a line approximating the gluteus mcdius muscle tendon.
In this example, a value for A of two times body
weight was derived from thc static free-body diagram
and confil'med by instll.llnentcd in vivo measurements (English & Kilvington, 1979; Rydell, 1966). The
direction of force A is found from a roentgenogrnrn to
be 30 from the vertical. The hOI'izontal and vcrtical
components of this force are found by vector analysis
(Calculation Box Fig. 8-2-3). The hOl-izonlal component (AJ is equal to body" weight; the vertical component (A y ) is approximately 1.7 times body \veight.
Attention is then directed to the lower free-body
(Calculation Box Fig 8-3-1). The gravitational forces
(Wand II. W) are known. The joint reaction force
(force J) has an unknown magnitude and direction
but originates from the most narrow joint space in
the radiograph and musl pass through the estimated center of rotation in the femoral head. The
magnitude of force J is determined by finding the
horizontal and vertical force components and
adding them (Calculation Box Fig. 8-3-2).
The value of J is found by vector addition (Calculation Box Fig. 8-3~3), and its direction is measured
on the parallelogram of forces. The joint reaction
force on the femoral head in a single-leg stance with
the pelvis leveled in the horizontal plane is found to
be approximately 2.7 times body weight, and its direction is 69· fTom the horizontal (Calculation Box
Fig, 8-3-4).
A key factor influencing the magnitude of the
joint reaction force on the femoral head is the ratio
of the abductor muscle rorce lever ann (c) to the
gravitational force lever arm (b) (Calculation Box
0
I
I
L
Fig. 8-2-2). This is particularly important in pros
thetic replacements of the hip joint (Delp &
Maloney. 1993; Free & Dell', 1996; Lim et al .. 1999
Sutherland et 'II., 1999; Vasavada et 'II., 1994). The
center of motion can be altered by the prosthctic de
sign and the lever arm for the abductor muscles can
be slightly changed by slu-get)! techniqucs. A change
of the center of location of the hip joint can de
crease the abduction force by more than 40(1'0 and
thereby the generated abduclor moment b~} almos
50% (Dell' & Maloney, 1993), Figure 8-11 illustrate
the relationship of this ratio to the joint reaction
force. ;\ low ratio (i.e., a small muscle force leve
ann and a large gravitational force lever arm) yield
a greater joint reaction force than docs a high ratio
A short lever arm of the nbduc(ol- musclc force, a
in coxa valga (Fig. 8-4), rcsults in a sTllall ratio and
lInls a sornewhat e1cvated joint reaction force. Mo\'
ing thc greater lrochanter latcrally during lata! hip
replacement lowers the joint reaction force. as it in
creases the lever arm ratio by increasing the muscl
force lever arm (Free & Delp, 1996). Inserting a
prosthetic CLIp deeper in the acetabulum reduce
the gravitational force lever ann, thereby increasing
the ratio and decreasing the joint rcaction force. I
is difficult. howc\'cl: to change the lever an11 ratio in
such a way as to reduce the joint rcaction force sig
nificantly been use the curve formed from plouing
the ratios becomes asymptotic when the ratio of c to
b approaches 0.8 (Fig, 8-1 I).
DYNAMICS
The loads on the hip joint during dynamic activi
ties have been studied by several investigator
(Andriacchi ct aI., 1980; Draganich c{ al.. 1980
English & Kilvington. 1979: Rydell, 1966). Using a
force plate system and kinematic data for the nOI"
Illal hip, J. P. Paul, 1967. (Forces at the human hip
joint. Unpublished doctoral theses, University o
Chicago) examined the joint reaction force on th
femoral head in normal men and women during
gait and correlated the peak magnitudes with spe
cine muscle activity recorded electromyograph i
cally. In the men, two peak forces were produced
during the stance phase when the abductor mus
cles contracted to stabilize the pelvis. One peak o
approximately four times body weight occurred
just nfter heel strike, and a large peak of approxi
matel:v seven times body weight was. reached jus
before toe-ofr (Fig. 8-12A). During foot flat, th
joint reaction force decreased to approximately
~.
I
Free-Body Diagram of the Lower Extremity
:
In Calculation Box Figure 8~3·1, the supporting lower ex-
tremity is considered as a free-body, and the forces acting on the free-body are identified. A, abductor muscle
force; J, joint reaction force; 1/6 W, gravitational force of
the limb; W, ground reaction force; Q, center of rotation
Tana
Jy
of the hip joint.
c··
2.5W
Tan a
In Calculation Box Figure 8-3-2, the forces acting on
a
A
Jx
2.5
69
the lower free-body arE> divided into horizontal and vertical components. Because the body is in force equilibrium,
the sum of the forces in the horizontal direction must
equal zero and
50
must the forces in the vertical direc-
tion. The horizontal and vertical forces are added and the
magnitudes of J.: and J( are found from force equilibrium
Jx
W
Calculation Box Figure 8~3-3,
equations:
A,~-J,~O
A,
~~J.-
'/W"," W ,-
a
J
I.7W-J.-'/W
W~O
A,~W
2~7
W
.I.
A
J..., 1.7 W
J,. - W
~,~
/. W
From the Calculation
Box 82 A. - I.7W
Calculation Box Figure 8-3-4.
Addition of the horizontal and vertical components J..- and
J, is performed graphically in Figure 8-3-3, and the joint
reaction force (J) is scaled off. A parallelogram is constructed and the diagonal of the parallelogram indicates
the inclination of the force J. Its inclination in relation to
the horizontal plane is measured (n) on the parallelogram,
Alternately, trigonometry is used to find the direction of J
using tangent equations (Fig. 8-3-4).
The joint reaction force has a magnitude of approximately 2.7 times body weight and acts at an angle 69"
from the horizontal,
w
Calculation Box
Figure 8~3~1.
w
Calculation Box
Figure 8-3-2.
Men
7.0
7
6.0
6
'la>
'ffi
:ca>
'ffi
;;
>1S
5.0
~
>-
"e0
~
~
0
5
4
@. 3
~
~
4.0
a
u.
2
u.
3.0
>
Upper and
lower
limits of
10<
angle of
2.0 ~..,._.,......,_..,.._.--..,._.,.......;5,,0:..'.., inclination
o
0:2
0:4
0:6
0.8 i of abductor
!1Bm
I,
I
I
I
Ratio of c to b
100
60
Percentage of CyCle
Women
muscle
force line
_
The value of the ratio of the abductor muscle force lever
arm (e) to the gravitational force lever arm (b) is plotted
against the joint reaction force on the femoral head in
units of body weight. Because the line of application of
the abductor muscle force (its angle of inclination in the
frontal plane) has finite upper and lower limits (10 and
100
60
Percentage of Cycle
50°), the force envelope is plotted. The curve can be uti·
lized to determine the minimal force acting on the
femoral head during a one-leg stance if the ratio of ( to b
is known. AdiJpced from Frankel. VH. (960). In The Femoral
Muscle AClivity During Slance Phase of Gait
Gll/leus ma~,mus
Sem:teM'llOSUS
Semimem!)r,lllOSUS
Biceps tomor,s
Neck: Function, Fracture Mechanisms, Internal fixation. Springfield: Charles C. Thomas,
I ---------------8
"======
r
~~~~~~~F1'"''
Tensor [osciac
S,lrtOri.us _
Gr,lc,bs
Reclus femoris
Gluteus med,us
Gluleus minimus
IIi",,",
E~tellsors
1------!--
.\dduc!or magous
Adductor IOr1(Jus
Adductor brevis
Abductols
.-==:==~AddUCIOIS
100
body weight because of the rapid deceleration of
the bodv's
centcr of ~
gravilv.
Durin t>lT the swing~
.'
.'
phase, the joint reaction force was influenced by
contraction of the extcnsor muscles in decelerating the thigh. and the magnitude remained relatively low. approximately equal to body weight.
In the women, the force pattern was the same but
the magnitude was sorncwhat loweI~ reaching a maximum of only approximalely four times body weight
at late sfance phase (Fig. 8-128). The lower magnitude of thc joint reaction force in the women may
have been the result of several factors: a wider female
pelvis. a difference in rhe inclination of the fcmoral
neck-to-shaft angle, a difference in footwear. and differences in the general pattern of gait.
c
60
Percentage of Cycle
Hip joint reaction force in units of body weight during
walking, one gait cycle. The shaded area indicates variations among subjects. A. Force pattern for normal men.
8, Force pattern for normal women. Adapted from Paul, IP
(1967). Forces at the human hip joint. Unpublished doctoral the·
sis. University of Chicago. C, Muscle activity during stance
phase of gait. The first peak corresponds mainly for the ex-
tensor and abductor muscles. The last peak. is for the flexor
and adductor muscles. Adapted from the University of California. (1953). The pattern of muscular acrivity in rhe lower excrem
i!'/ during ,·valking. Univ Cal Prosthet Dev Res Rep, 2(25). 1--4/.
•
Walking (0.9 mtsec)
- - Ground reaction force
.........u". Force on prosthesis
1400
1200
1000
~
..e
0
lL
t\
800
600
400
200
,
4
A
Walking (1.3 m/sec)
- - - - Ground reaction force
............." .... Force on prosthesis
1800
1600
In \"ivo ml..'asurl..'ll1enls of lhe forces acting
instrumented hip joint prosthesis demonstra
lower joint reaction force or the femoral head
ing the stance plwse or gail compared with ex
mcnsuremcnts and calculations (Rydell, 1965
8-13;\). At a faster cadence. the forces •.\Cling o
prosthesis grcatl~' increased because of an inc
in muscle ac(ivi(~' (Fig. 8-138). At both cade
the magnitude
the forces during swing phas
arproxirnatel~: half that during stance phase.
l~lble 8·2 summarizes the typical peak joint
on the hip joinl load e."pressed as bod~1 weight
different studies and with differ·ent methods
pattern of loading for \valking is similar for all
ies, but the magnitude of joint peak load differ
ternal measurements gencraHy yield higher
lated peak force on the hip joilll while inslrum
implant in vivo measuremcnts ~'ield lower
forces. There are man)' reasons 1'01' the diffe
for example, the melhod and instn1l11cnwtio
normal hip "crsus the "abnormal" instrulllcnte
plum. the gait velocity, and age. Activities othe
\\·alking, such as stair ascending/descending,
loads of around 2.6 to 5.5 body wciglll Illea
or
1400
1200
~
,
'"e0
lL
1000
,ImI!DJI
800
600
;!-
Range of Typical Reported Peak Hip Joint
Forces From Selected Studies
400
or-,
,1
200
SlanceDphase
Activity
i
i
1
2
i
3
Reported
Peak Force
BW
Instrumentation
... _--_.,.._,._--
-_.,-.'.'..._-_._---_
Walking
Swing phase
2.7-4.3
Insirurnented
implants
Bergmann
1993, 19
2.7-3.6
Instrumented
implantS
Kolzar et a
1991
2.7
Instrumented
implants
English et
1979
1.8-3.3
Instrumented
implants
Rydell et a
1966
4.9-7.0
EIviG
/force plate
Paul, 1967
45-7.5
EIviG
fforce plate
Crowninshi
et al.. 19
5.0-8.0
ElvtG
fforce picHe
Rohrle et a
1984
2.2-2.8
accelerometers
van den Bo
et aI., t9
Time (seconds)
B
Forces on an instrumented hip prosthesis during walking.
The broken line represents the force on the prosthesis, and
the solid line represents the ground reaction force. A.
Walking
Walking speed 0.9 m per second. B, Walking speed 1.3 m
per second. An increase in muscle activity at the faster cadence resulted in higher forces on the prosthesis. Adapted
from Rydell, tV. (1965), Forces in tfle hip-joint. Parr If: Intravital
measurements. In R.M. Kenedi (EdJ. Biomechanics and Related
Reference
Sio-Engineering Topics (pp. 351-357). Oxford: Pergamon Press.
BW. body \'1elght; EMG. electromyography
~--------'--'---'-"---'------"'-----r
Fatigue Fracture of the Hip
'1\.", 64-year-old, very active retired man experienced a
·M femoral neck fracture after changing his training
,'regime to prepare for a marathon. The fracture was
~.
'.> classified as a faligue fracture caused
by overload of
-, the hip joint.
;~
of a femoral neck fracture allowed a subseq
determination of the forces acting on the imp
during activilies or daily living (Fig. 8-14) (Fra
el aI., 1971). Although the device measured fo
on the implant and not on the hip joint, it was
sible to determine the proportion of the
transmitted L1nough the device and to calcu
the total load acting on the hip joint by mean
static analysis. In the case illustrated in Fi
8-14. the nail plate transmitted one fOllrth of
lOtalload.
Strong forces aCling on the nail plate were
countered during such diverse activities as mo
onto a bedpan, transferring to a wheelchair,
walking. The magnitude of the forces was gr
modified by skillful assistance frol11 lhe nurs
therapist to control the patient's movement. Fo
of lip to fOllr limes boely weight acted on the
joint when the patient used the elbows and hee
elevate the hips while being placed on a bed
(Fig. 8-15), but these forces were greatly red
through the lise of a Lrapeze and assistance from
attendant (Fig. 8-158). A 5-kg ex.tension tractio
the hip had little effect in modifying the forces
ing on the hip joint. Exercises of the foot and a
increased these forces.
Case Study Figure 8-2-1.
The figure shows an MRI (frontal view) of the pelvis
and both hip joints. The fracture is seen in the left femoral
neck distal to the femoral head. The fracture is believed 10
have occurred during running and after an extensive
change of training program. Because of the high repetitive
I?ading, muscle fatigue, and the change in the load pattern, on the hip joint and femoral neck, the bone fractured.
with an instnlmentcd hip implant (Bergmann et al.,
1995; KOlzar et aL. 1991). The highest magnitudes of
load during daily activities are measured dllling
stair climbing and getting up from a low chair when
the hip is flexeell1lore than 100" (Catani et aI., 1995;
Johnston et aI., 1979). Co-contraction of the biarticular muscles was evident during these activities.
Running and skiing lIsing accelerometers yielded
calculated forces up to eight limes body weight in
middle-aged and older people (van den Bogert et al.,
1999) (Case Stlldy 8-2).
Insertion of an instrumented nail plate in the
p'-oximal femur arter osteotomy or during fixation
..,
...,:
..
An instrumented nail plate in the proximal end of the
mur was used to determine the forces acting on the i
plant during the activities of daily living following fra
of the femoral neck. In this ca5e, the nail plate was fo
to transmit one fourth of the total load on the hip jo
Usc of the instrumented nail plate demonstr
that, for a bedridden patient with a fractu
femoral neck, the forces on the fernoral head
ing activities of daily' living approached those
iog walking with external supports. These stu
support clinical protocols for carly mobilizatio
patients and decreased bed rest for patients
A
B
A, When the patient used elbows and heels to elevate the
hips while being placed on a bedpan, the force on the tip
of the instrumented nail was 670 N. With a spica cast, the
force on the tip of the nail was 190 N. B, The use of a
trapeze and assistance from an attendant reduced the
force to 190 N without a cast and to 70 N with a spica cast.
Reprinted "vith permission from Frankel, VH. (7973). Biomechanics of the hip. In R. G. Tronzo (feU Surgery of the Hip Joint
(.0,0. 105-125). Philadelphia: Lea & Febiger
hip fractures. The magnitude (approximatel
Nlll) of the moments acling on the nail-plate .
lion in the transverse plane (i.e .. during inte
and extcrn~ll rotation) was only approxirnatel)!
half the magnitude (18 Nm) of the moments ac
in the frontal plane (Le., during abduction)
many activities.
EFFECT OF EXTERNAL SUPPORT ON THE
HIP JOINT REACTION FORCE
Static anal.\·sis of the joint reaction force on
femoral head during walking with a cane dem
strates that the cane should be used on the side
posite the painful or operated hip. Neumann (1
High and low load on the hip joint during daily activities. Raising from a low chair produces
approximately 8 times body weight (A). Walking with a cane on the ipsilateral side of the
affected hip produces approximately 3.4 times body weight (B), and walking with a cane on
the contralateral side of the affected hip reduces hip joint load substantially to 2.2 times
body weight (C). This figure illustrates how load on the hip joint can be manipulated by
simple means (X denotes affected hip).
-:i:::~C::I::J::r:S:~ t:~::r::e~h:i::~re:~~o<~Oint Rea~:~::e:::C:uSCle
I
faKe acting on the femoral head in Ihe late swing phase of
I the
gait cycle for an 8·year-old boy INeighing 24 kg and
momenl relationship
T ::::: Fd,
i
I
where F is the extensor muscle force and d is the perpendicu~
lar distance from the center of rotation of the femur to the
wearing a )ong-Ieg brace. The main muscle force was produced by contraction of the gluteus maximus muscle and
identified through electromyography. The torque about the
middle of the gluteus maxim us muscle. Distance d was mea.,
hip joint \ivas calculated according to the formula
sureej from a roentgenogram and found to be 3.2 em, From
H'e equation E;::;: TId, the muscle force on the normal Side',
. . " as calculated to be 338 N, and on the braced side, 600 N.
T
~
la,
where
T
is the torque expressed in newton meters (Nm)
I
is the
The joint reanion force on the femoral head (J) is equal to
mass moment of inertia expressed in newton meters
times seconds squared (Nm sec!)
is the angular acceleration in late swing phase, expressed
-;,-
in radians per second squared
(rlsec 1).
In the case of (he braced side. I = I , .;. I;; where
is the mass moment of inertia of the leg
is the mass moment of inertia of the brace.
On the normal side,
I "" 0.45 Nm se('
n=:
24 rlsec'.
Thus,
I = 0.45 Nm sec:'
.: ~
force (E) was then found from the
x 24
T
~
On the braced side.
1 0.45 Nm sec; + 0.35 Ntn sec'
mated to be 40 N.
On the normal side,
On the braced side,
J = E -\iV,
J = 338 N - 40 N
J = 298 N.
J = E -\iV,
J=600N-40N
J = 560 N.
Thus, tile Joint reaction force on the femoral head in the
br~1Ced limb v"as over 80% higher than the force in the non~
;.cc
n =
24 rlsee"
braced limb, reaching more than two times body weight.
Thus,
T = (.45 Nm sec) + 35 Nm sec!)
x 24
rlsec~
10.8 Nm.
the muscle force (E) minus the gravitational force produced
by Ihe weight of the limb (W,). In this example, W L \·\las esti-
T
~
rlsec~
19.2 Nm.
I
0~---
,;::
studied the effects of cane use in 24 subjects with a
mean age of 63 years. During walking, the e1ectromyographic activity of the hip abductor muscles
was measured, Neumann found that usc of a cane
on the contralateral side of the affected hip joint,
with careful instructions to use with near maximal
effort, could reduce the muscle activity by 42% (Fig.
8-16). This calculates to a reduction of approximately one times bod~; \veight from 2.2 body weight
with a cane, compared with 3.4 body weight withOut a cane. These studies give important inforllla~
lion to the clinician about ways to moderate the
load for the patient with hip problems.
Such use reduces the force on the femoral head
of the painf-ul joint without necessitating an anlalgic
body position. A cane used on the side of the painful
hip works through a shoner levcr arm and thus an
even greater push on the cane is needed to decrease
the joint reaction force. For the olcler patient, such
a large push may not be possible because of a
of strength in the upper extremities.
The use of a brace on the leg may aller the f
on the hip joint but may not always reduce the
reaction force on the femoral head. An ischial
leg brace used in the treatment of Perthes' dis
raises the joint reaction force during late swing p
because the large mass moment of inert ia o
brace results in a higher extensor muscle force
ing this part of the gait cycle (Calculation Box 8
SUl1lmary
1'; The hip JOll1t is a ball-and-socket joint
posed of the acetabulum and femoral head,
2 The thickness and mechanical properties o
on the femoral head and acetabulum
from point to poinl.
cartilage
-_.~,
3 Hip llexion of at lca~·a 120", abduction of at
least 20", and extenwl rOl<.l.lion of at least 20" are
necessary for can}'ing out daily activities in a normal manner.
\4 A joint reaction force of apprOXill'latc1y three
timcs body weight acts on the hip joint during a single-leg stance with the pelvis in a neutral position;
its m;gnitudc varies as the position of the upper
body changes.
_~'> The magnitude of the hip joint reaction force is
infiuenced by the ratio of Ihe abductor muscle force
and gravitational force lever anns. A low ratio yields
a greater joint reaction force than does a high ratio_
;;_~ The hip ,joint reaction force during gail reaches
levels of three to six limes body weight or more in
stance phase and is approximately equal to body
weight during swing phase.
j.7 An incrense in gait velocity increases the magnitude of the hip joint reaction force in both swing
and stance phase.
,'8 The forces acting on an internal fixation device
during the ~\ctivitie; of daily living vary greatly'
depending on the nursing carc and the therapeutic
activities undertaken by the patient.
/9 The lise of a cane or a brace on the leg can altcrthe magnitude of the hip joint reaction force.
REFERENCES
Andriacl:hi. T.P.. Andersson. G.B .. Famil:r. R.W., Stern. D..
G,t!anil', J.O. (1980). r\ study of to\\'er-limb medlanics during stair·dimhing. J BOllI.' Johu Surg, 62.-1. 749.
Bergmann, G., Graichl'll. E, & Rohlm=lIl1l. A. (1993). Hip joint
1~'H.ling during \V,lIking i.lIld running llh..· .ISurcd in IWo pa·
tknts. 1 Bio/lw:h, 26(8),969-990.
Bergmann. G.. Graichen. E. & Rohlmann, A. (t995). ts slair..
cast: walking a risk for tht: fi.X<lli(lll of hip implants?} 13ioIlIe,''', 28(5), 535-553.
Catani, F.. Hod!!(', r\ .. ~'talln. R.W.. et ::d. (1995). The 1"011.; of
museul"r co:eontr<lClion of the hip during mon.·men!. Chi,.
Orgalli ;\1ov, 80(2), 227-236.
Crowllinshidd, R.D .. Johnston, R.C.. Br<lnd. R.t\. (1978). The
eHccts of walking no'locity and agl' on hip kint:m<ltics and
kinetics. Clill Or/hop Rd Ues, 132,140-14-\.
Ddp, S.L. &. Maloney, \V. (1993l. Effects of hip Centcr locatioll on the momcnl-generating cilpacity of the l1luscks.)
8'0mi'd,. 26(5). 485-499.
Drng::lllil·h. L.F.. Andrincchi. T.P.. Strongw::llcr. :\.1\'1.. ct ~\1.
(l980). Electronic tlll.:aSLlrcrnent
instantaneous fOOLfloor conl;:\ct paltl'rns during gait. 1 BiOlIICdl. 13. S75.
English. T.A. & Kilvington, M. (1979). In \'ivo re('ord~ of hip
IO:.lds using ::1 fl'llloral impblll with lelemetric output (a
or
preliminary l't:ponl. 1 Biollled ElIg, 1(2), 111.
Fl·ss.\', ~1.I-1.. ;\'Diaye, A.. CalTd. J.P.. et :II. (1999)_ Lo
tht..· l·{.·lltl-r or rOlation of IIII.' hip. Sl/r.g nmliol A//at,
24;-250.
FI'~lnkd, V.H. (19;3). Biollll'challil:s
the hip. In R.G. T
(Ed.), Slll'.~t:ry o( I!lt' Hip }Oil/l (pp. 105-125). Phil~ldl.
lea &. Fdlif!l'r.
Frankd. V.H. (1960). In nUt FI.'J/ulI"iI} Nt'd: {-'Wlclioll. F
.\-1f.'(:IUllli:'/II.\, "/lenla! FixaTioll. Springfidd: Char
or
TlwllWS.
Fr'lnkd. V.H .. BurSil'in. A.H .. L~·grl·. L., l't al. (1971). Th
L.lil: nail. ) BOIH' 10illl SIII"~. 53:1. 1232.
Free, S.A. & Oelp, S.L. (1996). Trochantt.'rll.: transfer i
hip n:pJal:i.:IlIl.:nl: Effecls on Ilit' lTlOfllL'llt arms nnd
gellt..'ralillg c'lpadtic:;s of Ihe hip abductors. J Onlw
1-I{2L 1-15-250.
Gn.·em\'~a1d.
A.S. & 1-I<I\'llI,:S, D.W. (1972). Weight-hcaring
ill lhe human hip j{~Jinl..! BOIlt' .!Oilll Siltg. 5-1B( I), 15
Hurwitz. D.E. & Andri.tt:("hi. '1'.1>. (1998). Biomechanics
hip. III J.J. Cdbgh'111. A.G. Rosl'nberg. &. I-I.E. R
(Eds.1. The: Adult flip (pp. 73-85). Philaddphia: Lipp
Raven Publishers.
Hllr",it;',. D.E. & Andri~\(:("hi. T.P. (1997). Biolllcl:hanks
hip .tlld the knl'l'. In ~'1. Nordin, G.B.J .•\ndnsson. &
POPl' (Eds.) . .\lusculoskclewl Disorders ill {Itt' Irork
Principles (Jud Prauice (pp. -186--196). Phila(k']phia: M
Year Book.
Inrn;lIl. V.T. (19-47). Flll1l:lional aspects of t!ll' abductO
c1es of thl' hip. J BOllI.' loim Sll"~. 29:\. 607.
JohnslOn, R.C., Br~Hl(I, R.A .. &. Crowninshicld. R.D. (
Re('onslnH:lioll of tht' hip. } BOI/t' Join, Smg. 6
646-652.
Johnston. R.C. & Smidt. C.L. (1969). !\leasur(,llll'nt o
join! motion during walking. Evaluation of 'tll I'll'
niOllll'tric Illt.'lhod. 1 BOll...' loiJ/! Sll}".~. 51:\, 1083.
Johnston, R.C. & Smidl. C.L. (1970). Hip motion tlll
m('Il{S for sell-tted .\Ctidties of daily li,·ing. efill O
72. 105.
Kcmpson, G.E., Spin.'y. c.J .. Swanson. S.A.V.. cl at. (
Patk'rns of C<lrtila~1' stiffnt..'ss on normal and dC~Cl
human femoral he;ds, J liiO/IIt'ch, -I, 59;.
Kim, Y.T. &: Azurll.l, H. (1995). The lH.'I'\'e t,;ndings of
etabular "tbrum. CfiJl Or/hop, 320, 1;6-181.
Konrath, G.A., Hamel. A.1 .. Olson, S.r\ .. d .d. (19981. T
of the aCI'I.dHlbr labrum and the (r~\IlS,'e..SI' :lcL'(ablll
;l11lt.:lIt in load transmission of the !lip. J (JOllt~ loill
80A(I2I. t781-1788.
Kotzar. G.~L Davy. O.T.. Goldberg. V.;\1., cI al. (
TcJemilriZL'd in "j\'o hip joint forcc dilta. A report o
ll<ltit.:IHS aftcr 10iai hip sllrgery. 1 Ort/wp Nt's, 9. 621
Kumagai. iv1.. Sliiba, N., Higuchi, F.. et a!. (19971. FUll
t.:\'aIU:llion of hip abductor mllSC!<;S with llSI' of m.
resonance imaging. 1 Orlhop Rt,S, 15(6),888-893.
lilli. l.A .. CIl"mi(;h;lel, S.W.. &. C<lbancla, M.E. (19991
llH:chanics of lOUt! hip ;lrlhroplnsly. AI/ot UI:C, 2
t 10-t 16.
j\'kLl'ish, R.D. & Chilrnh:y. J. (1970). Abduuion forct's
one-!l.'ggcd st~lnc('. J B;oll/cch. 3, 191.
;\'ld,linn, ~R.H. &: Huchil1!.;s, R.H.R. (1988). In C%}" Il
HI/lllml AWIlO11/Y (2Il1.! cd., p. 302). Chic,lgO: )\:.11
,\-kdic<ll Publishl'rs, Inc
MlllT~I\'. ~LP. (1967).. G.dt <lS ,I lOt~t1 pattern of moveme
1 pjlYs .lled, -16, 290.
\-!urray, i'",1.P., Kor.\', R.C., &: Clarkson, 13.11. (1969). Walking
patterns in he:dthy old nH:n. } Geroll/ol, 2-1, 169-178.
Nemeth, G. &: Ohlsen, I-l. (1985). In vivo moment arm lengths
for hip extensor muscles at difTerent angles of hip fle\ion.
} Biolllech, 18, 129-149.
Nemeth, G. & Ohlsen, H. (1989). rvlolTlel1t arms or hip abductor and adductor muscles in vivo by computed tomogra·
phy. Cfill 13iolllccll, 4, 133~136.
Neumann, D,A, (1998). Hip abductor JllLlscle activity as subjects with hip prostheses walk with dirrerL~nt methods of
'using a cane. Pllys Ther, 78(5),490-501.
IWhrle, H., Scholten R., Sigolo!to, C., et al. (1984). Joint
forces in the human pelvis-leg skeleton during walking. }
Biolllcch, 17, 409-424.
Rushfcld, P.D., \tann, R.\V., &: Harris, W.H. (1979). Influence
of cartilage geometry on the pressure distribution in the
human hip joint. Sciellce, April 27, 204(4391), 413-415.
Rydell, N.W. (1966). Forces acting on the remoral head prosthesis: A study on strain gauge supplied prostheses in living persons. Acta Orr/lOp S("(lIld, SlIppl88. 1-132.
Rvdell, N. (1965). Forces in the hip-joint. Part If: Intravital
. measurements. In R.rvl. Kenedi (i~d.), l3iolllt'cllilUics (jlld
Bio-LII~il/ecrilig Topics (pp. 351-357). Ox
f\:rgamon Press.
Sutherland, A.G., D'Arcy, S .. Srn:lrt, D., et al. (1999). Ab
tor muscle weakness and stress around acetabu!:lr co
nents of total hip arthroplasty: A finite dement ana
Re!uh'd
1111
Orthop. 23(5), 275-278.
University of California. (1953), The pattern of muscula
tivity in tht: lo\\'er extremity during walking, (jllll
Pros/hd DcI' Rcs Rep, 2(25). 1-41.
van den Bogert, A.J., Read, L., &. Nigg, B.rv1. (1999). An ;
sis of hip joint loading during walking, running and
ing . .lIed Sci Sports, 31(l), 131-142.
Vas:lvada, A.N., Delp, S.L., \.laloney. \V.J., et a!' (1994).
pens;:lling (or changes in Illuscle length in total hip <l
pbst~,. Effects on tllL' llHHTll'nl gt:nt:rating capacity o
muscles. Cfill OnllOp. 302. 121-·133.
VonF.isenhart-Rotht:, R.. Eckstein, r.. ~luller-Gerbl. rlit.,
(1997). Direct comparison or contact <.In:as, contact
and subchondral mineralization in Illunan hip joint s
mens . .. \I/IJ{ f:'lIlbryol ([kr/), 195(3), 279-288.
Biomechanics of the
Foot and Ankle
G, James Sammarco, Ross Todd Hockenbu
Introduction
Growth of the Foot
Kinematics of the Foot
Foot and Ankle Motion During Gait
Causes of leg Rotation During the Gait Cyde
Muscle Action During Gait
MotIon of the Tarsal Bones
Sublalar Joint Motion
Transverse Tarsal Joint Motion
Tarsometatarsal and Intertarsal Motion
Motion of the Hallux
Motion of the Lesser Toes
The Medial longitudinal Arch
Muscle Control of the Foot
Kinetics of the Foot
Soft Tissues of the Foot
Ankle Joint Biomechanics
Kinemaiics
Range of Motion
Surface Joint Motion
Ankle Joint Stability
Kinetics of the Ankle Joint
Statics
Ankle load Distribution
Dynamics
Effects of Shoewear on Foot/Ankle Biomechanics
Summary
References
Introduction
The biomechanics of the fool and ankle arc complex
and intricately associated with each olher. The fOOl
is an integral mechanical part of the lower cxtrcrnity necessary 1'01' a smooth and stable gaiL The ankle transfers load frolll the lower extremity to the
foal and closely inllucnccs foot orientation with the
ground.
The fOOL is comprised of 28 bones (including
scsamoids) whose motions arc closely inlcn-c1alcd
(Fig. 9-1). Besides acting as a structural sllpponing
platform capable of withstanding repetitive loads or
mulliples of body weighl. the fOOL/ankle complex
also must be able to adjust to different ground surfaces and vat)'ing speeds of IOCOl11CHion. The unique
qualities of the 1'001 allow it to be rigid when ncccsSal)', as in ballet dancing on point, or quile flexible,
as in walking barefoot on the sand. The transition
from shock-absorbing platform to rigid lever capable of forward propulsion occurs wilh cach slep of
lhe gail cycle.
The ankle is compriscd of lhrec boncs Ihal form
the ankle mortise. This joint complex consists of the
libiotalar, flbulotalar, and tibiofibular joints (Fig.
9-2). The ankle is a hinge joint whose stability depends on joint congruency and the medial. lateral,
and syndesmotic ligaments.
This chapter discusses the motions that occur in
the foot and ankle during the various phases of gait
as well as during extremes or motion. The close interplay between lower extremity rotation and Forefoot orientation is explained. The ground (foot-tofloor) reaclion force and dislribulion of forccs on
lhe planlar aspeci of the fOOL are explored. The 10calion of forces as they pass from the tibiofibular
complex into the dome of the talus and then into the
foot is disclissed. Vve also discuss the roles of ligaments and muscles in the support of the medial longitudinal arch. Finally, ankle motion and ligamentous stability is outlined.
A discussion of sophislicated electromyographic
activity during walking is not within lhe scope of
this chapter; however, the activity of certain extrinsic and intrinsic muscles is by necessity presented to
allow a beller understanding of foot and ankle control during gait. The moments produced about
joints by muscle action and resultant effects on foot
and ankle position are detailed. Joint axes and instant centers of joint motion are described. Refer Lo
Chapler 18 for information aboul the application of
biomechanics to gail.
Any pathological change in root or ankle structure or motion, however subtle, may have a pro-
found impact on the foot and ankle's shock-abs
ing, propulsive, and stabiliZing roles. Clinical c
lation of alterations in biomechanical Functio
presented in case studies. Footwear in \Vestern
ciety may VaJ'y from a rigid ski boot to a soft
casin. These externally restrictive materials ma
ter normal foot and ankle biomechanics
ultimately innucncc the development or s
pathological conditions, slich as hallux valgus.
Growth
of the
Foot
The foot is formcd when the limb buds develop
ing the eighth week of gestation. Foot length
width increases linearly from age 3 to 12 in
and age 3 to 15 in boys at an average of 8 to 10
per year, followed by a plateau in growth (Che
al.. 1997). Blais and associales (1956) showcd
the foot appe«rs to be closer to the adult size a
times during normal development of the child
arc other parts or the limb. On average, at a
)'ear in girls and 18 months in boys, the leng
the foot is one half the length of the respective a
foot (Fig. 9-3). This situation contrasts with th
the femur and tibia, which do not attain their
ture length unlil 3 years later in both boys and g
The relativel.\' large size or the fool, then, is im
tant for providing a broad base on which the ch
body is supported. and Lhis base may at tilTlCS c
pensate for the child's lack of muscle strength
coordination.
Kinematics of the Foot
Gross motion of the foot is complex and oc
m'ound three axes and on three planes (Fig.
Flexion-extension occurs in Lhe sagittal p
abduction-adduction occurs in Lhe horizonta
transverse plane, and inversion-eversion occu
the coronal or frontal plane. Supination
pronation are terms commonly used to desc
positioning of the plantar surface of the foot
occur primarily at the subtalar (talocalcan
joint. During supination the sale faces medi
and during pronation the sole faces lnter
Supination is a combination of inversion, flex
and adduction. Pronation is a combinatio
eversion, extension, and abduction (Fig. 9-5).
motion includes Oexion, extension: adduction.
abduclion.
For practical purposes. foot motion can be
sidered to be of two distinct types; non-we
Base\
"
\ 'KJ'"'R,..
Os trigonum
Calcaneus
Head
Phalanges
\
Sesamoids
Tuberosity
MEDIAL VIEW
Trochlea
Os
Metatarsals
trigonum
Tuberosity
Point or insenion
01 the carcaneolibular
ligament
calcan~-~~"""""'''~
-.;...·..><'·:hlea ('ubercle)
/
Groove lor
.
TuberOSIty
peroneus longus
Head
TU!)fHc!e
LATERAL VIEW
Calcaneus
Latera! process
Talus -"'- ~~,.
Medial
malleolus
lateral
Cuneilorms
{
Middle
Lateral malleolus
Medial
r"rsomelalarsal joint
Tibiofibular articulation
(of Lisfranc)
HaUul(
-Great 100-
DORSAL VIEW
Top, View of the medial aspect of the foot. Middle. View of the lateral aspect of the foot.
Bottom left, Superior view of the foot. Bottom right, Anterior view of the ankle mortise,
Med.
La!.
Interosseous
membrane
L
Fibula
t
. . ~istal
llbloflbular
•t
-~-'-......:
.<
z:f,
....z..
0
0~,
Ii I
(~¥J~Tjbia
f&t~~:
;:;if4~.
...-;&"4.~
::~.,-:"~r-7: •.•
Tibiotalar joint
II
~
/
:R'f1oJ-~-
Calcaneus
ft
~,
~
f
I
Ankle joint complex composed of the tibiotalar, fibulotalar.
and distal tibiofibular joints.
f
beal"ing and weight-bearing. Passive. nOI1-wdgh
bearing Illotion may be tested with the patien
seated and the foot and ankle hanging free. Subtala
motion is evaluated by grasping the tibia with on
hand and inverting and cvcrting the heel with th
other hane!. Abduction and adduction of the fore
foot can be tested if the heel is held immobile
Supination and pronation of the forefoot may als
be tested with the heel fixed, as may flexion and ex
tension of the tarsometatarsal joints and [Des.
Active, weight-hearing 1TI00ion of the foot differ
from passive motion because the forces produced b
body weight and by muscle contraction act to stab
lize the joints. Generally, functional active foot mo
tion during gait tends to be less than passive foot mo
tion. Active subtalar inversion can be demonstrate
by viewing heel orientation from behind while askin
the patient to raise lip on his 01' her toes. Extel'nal ro
tation of the leg while bearing weight on the foo
causes the hed to invert and the forefoot to pronate
therefore raising rhe arch. Rotating the leg internall
has rhe opposite effect: it lowers the arch.
~'
F
~;:
g
i
t
f
f
I~
;'
~
I
i
~
F
~'
LENGTH OF NORMAL FOOT
Girls
Boys
30 ,----,----,----,----,----,----,----,----,----,
28
+-+-+-f-+-+-+-+-+--I
.1. !.!
26 +-+--+--+-+-+-++'-+--1
V
24\
·
E 22 W-+-~~·~·····
30 ~
28
I I I I I I I I.! . l
I Iii ' I
26
+1-+1--h-t-:·,pv';l"/~····--4
I .. i<;V i
I I·l/;{.
§ 18:/
§ i
lOX)· i
16 t-+--flfC'fJ+'-I-+--T.:-'-...L-j 18 \
all· I I
1,0
P.'cenmes
14l-l!f+-++-+--1·······..· · · ·
/ -
1 +1
24+-
4+=1EE 22\
./...' ::.:& 20 I
::.~ 20
0
50
0
I
16\ i/.'./I/!·
'I
14 +-fff!l+-+-+-+--l---'f---H
I
\
10'O±jjt±±t::±j lot,±l=l:±jj=±jj
... ~59~5
.f.
12
o
·i
2
4 6 8 10 12 14 16 18
Age (years)
%' • I
12 H'l-+-+-+-+--l-f---H
0
2 4
6 8 10 12 14 1618
Age (years)
I.
lengths of normal girls' and boys' feet derived from serial measures of 512 children aged 1
to 18 years. Left, Foot length versus age in girls. Note plateau in growth after age 12.
Right, Foot length versus age in boys. Note plateau in growth after age 15. Adapted with
permission from Blais. M.M.• Green. W T.. & Anderson. M. (1956). Lengths of the growing (oot. J
Bone Joint Surg. 38A. 988.
thlllst on the subwlar joill!. In the middle or
phase and at push-ofr. the entire lower cxtrcll
Dorsiflexionl
planlarflexion
Eversion!
inversion
~i~lS t(~ reverse and ,:otatc L·.'(t:nlall~' as ..1he. ~
JOint Sll1lllltancollsl~' II1vcrts (Fig. 9-9). \\illh IIlV
of the subtalar joint and supination 01" the fo
foot is transformed into a rigid stntclurc cap
propulsion. Olerud and Rosendahl (1987) and
berg Cl al. (1989a-d) Im"e experiml'ntallv me
lhe coupling 01 tibial rot~lliOIl to subtalar m
They h<:\ve shown that lhe foot supimues I" fo
0.2 to 0.44" 01" libial external rotation.
Causes of Leg Rotation During the Gait Cycl
Malln (1993) has described the coordination o
Foot motion occurs around three axes.
~~--------
and subtalnr rl'lotion in an ekg::HH model c
mitered hingL" (Fig. 9-10). As the tibia interna
lales. the subtalar joint evens (pronatcs). COl1
external rotation or the tibia causes in
(supination) of the subtalar joint. He attribu
internal rotation or th(: lower cx.t1\:l11it~: in
stance to the obliquit.\' 01" a general ankk joi
(see Ankle Joint l\'totion). According [0 the "
axis model
the ankle joint, the ankle joint
angkd downward and postl.'riorly I'rorn Ilk·dia
eral (Fig. 9-1 I). Betau:-ic of the obUquily of th
or
FOOT AND ANKLE MOTION DURING GAIT
The gait cycle consists or a stance phase and a swing
phase. The stance phase encompasses 62(1'0 of the
gait cycle and the swing phase makes up the remaining 38%. The stance phase is subdivided into
heel strike, foot flat, heel rise, push-orr. and toe-off.
Swing phase is divided into acceleration, toe clearance, and deceleration phases (Fig. 9-6). The part or
stance phase spent with both feet on the ground is
termed double limb support and occurs through the
flrst and last 121'10 of stance phase (Fig. 9-7). N0l111al
men have an average gait velocity of 82 m/min and 58
heel strikes/min (Waters et aI., 1978). Running is defined as a gait speed pasl 201 m/min. At this speed,
double stance disappears and a noat phase develops
in which bOlh reet are orr or the ground (Fig. 9-8).
During normal walking, the entire lower ex.tremity
(including the pelvis, femUl~ and tibia) rotates internally through the first 15% of stance phase. From
heel strike through foot flat the subtalar joint everts,
tbe foot pronates. and the rorefoot becomes flexible
to absorb shock and adapt to irregularities in the·
ground 11001' surface. The subtalar joint everts in part
because the point or contact or the heel is lateral to
the center of the ankle joint, thus producing a valgus
A, During foot supination the sole faces medially.
ing foot pronation the sole faces laterally.
•
I
I Heel stnke
~=K\~ 7TTcr 5l~
ll~l/~ 1~lf~1
I
! I Heel slnke
I L 0°"
I
/1
:
Fool flat
I
15~o
l~~l-p-"-Sh-O-lI~'-T-o-e-o-I-,-1
30~'o
I
45~'o
I
1
IAcceleration [Toe c1em""e I Decelemtion
_~~~~
Stance phase
Swing phase
<l~--------_~
.~----_~
60~o
Double
Double
limb
support
limb
Single limb
support
~
<1-+
support
~
<1-+
I
SWing Phase
70%
Heel st
II////II!
Stance Phase
~I2J~]~~
I
Toe ot!
iH~e~
100~,,
l
i
IIBII
..
62
Percent of Walk Cycle
'--~~~~~
_
Stance phase consists of two periods of double limb su
port and one period of single limb support.
62% of the normal gait cycle is spent in stance phase and
38% is spent in swing phase.
I
&------------------
A
Stance ~65'
:
o
10
Right heel
Q
RighI twel
slrike
;
If:
20
30
:'
II
40
,
50
:
j
SWina(?5~o)
70
60
Mid
I. osl"rt~k_e
A
AA A
Ji
Walking
80
90
100
Left heel
,-st-a.L:c-eT--T--,-,-F-O-'t-O-'-t'---S-1lr~ke
•
Mid
slance
+ +
~
+
'I''I'
Toe Double limb ~:~tl Mid Double limb Right heel
off unsupported strike stance unsupported slnke
Running
Stance (40%)
o
10
20
30
Swing (30'%)
40
50
60
70
80
90
100
{ ~ftJz~~~~
I
i
IBDI'-i
•
1
Comparison of walking and running cycles. In the running cycle, stance phase decreases,
9p
float
develops: _d_O_U_b_l_e_l_im_b_S_U_P_P_O_rl_d_iS_a_p_p_e_a_rs_,_a_n_d_a_d_o_U_b_l_e_l_im_b_U_n_s_u_p_p_o_r_te_d_o_r
S_W_in__phase
_h_a_se_in_c_r_e_a_se_S_'
__
_
..
..
..
Toe-off Heel contact
Heel contacl
Ankle _.
rotation
i
!
'
!.......-:-- Stance phase
,6~T
.:t;
:....
_:
,I
FOOl immobile ....:
Subia la, ~_.'-_..__
rolation
:-
B
A
f,'
I
°
,._L_~_~.. -----,----;c;;~'
m
~
I
-----'_._~
W
W
I
I
100
I
Percent of Walking Cycle
I
I
l
Ankle motion and subtalar rotation during normal walking.
Maximal subtalar eversion occurs at foot flat in early stance
phase. Maximum subtalar inversion occurs at toe-off.
•
joint axis, the leg internally rotates with ankle dorsi~
flexion and externally rotates with ankle plal1larflexion. Additional mechanisms by which leg external rotation occurs during late stance include the swing of
the opposite leg that causes external rotation of the
planted leg and the obliquity of the metatarsal break
(Fig. 9- I 2). The metatarsal break is an oblique axis of
50 to 70" with respect to the long axis of the [oat
formed by the centers of rotation of the metatarsophalangeal joints. With push-off. the foot and
lower extremity externally rotate with respect to the
sagittal plane because of this oblique axis.
c
o
Mitered hinge model of leg, ankle, and subtalar mo
A. Outward rotation of the upper stick causes inwa
tion of the lower stick. 8. Inward rotation of the up
stick causes outward rotation of the lower stick. C.
nal tibial rotation causes supination of the foot. O.
tibial rotation causes pronation of the foot.
Number of
specimens
MUSCLE ACTION DURING GAIT
20
Although the motions of the foot and ankle during
the walking cycle OCClIr primarily as a result of the
passive constraints of joints and ligaments. electrom.yography has shown that rnuscle activity does
occur during normal gait (Fig. 9-13). At heel strike,
the pretibial musculature fires eccentrically to slow
down the descent of the forefoot and prevent a [oat
slap. At midstance, the calf musculature contracls
to slow down the Fonvard movement of the body
over the foot and prevent a crouch gait. The intrinsics also contract during midstance to toe-off to aid
in rigidity of the forefoot. Toe-off is primarily a passive event. The pretibial musculature again contracts during swing phase to ensure (hat the fOOl
clears the Ooor during swing~lhrough.
15 r
r
I
10'
{
74 78 82 86 90 94
Angle (degrees)
Variations in the angle between the midline of the
and the empirical axis of the ankle. The axis is angle
obliquely and inferoJaterally 82°. The histogram sho
variability among specimens.
Metatarsal
break
version and eversion. The sublalar joint is respo
sible along with the transverse tarsal joint (consis
ing of the talonavicular and calcaneocuboid joint
for transforming tibial rotation into forefo
supination and pronation. Because the ankle Joi
is to some degree a single-axis joint. subtalar m
tion reduces the rOLalol y ' stresses to the ankle join
Congenitally blocked motion of the subtalar joi
may result in the formation of a ball-aod-sock
ankle as a result of the increased rota1011' stress
the joint. Manter (1941) determined Ihe subtal
axis of rotation to be oriented upward at an ang
of 42" fmm the horizontal and medially 16" fro
the midline (Fig. 9-14). The sublalar facets resem
ble segments of a "spiral of Archimedes," a righ
handed screw in the right foot, so that the calc
neus actually translates anteriorly along t
subralar axis as it rotates clockwise during the m
tion of subtalar inversion (Fig. 9-15). A\·erage su
talar motion is 20 to 30° inversion and 5 to 1
eversion. Functional subtalar joint motion duri
gait is 10 to 15°. During the gait cycle, the he
strikes the ground in slight inversion, followed
rapid eversion to a maXimUI11 of 5 to 10° at 10%
the gait cycle (Fig. 9-9) (Sarrafian, 1993a,b).
Transverse Tarsal Joint Motion
The transverse tarsal joint, Chopart's joint, consists
the talonavicular joint and calcaneocuboid joi
Manier (1941) described two axes of motion in t
I
I
Muscle activity I
$
,;.:
The muscles of the lower extremity are more aclive dudng nmning. The gluteus maximus and hamstrings arc active in midstance through toe-off and
increase their activity 30 to 500/0 1O decelerate the
stance phase limb. Dorsillexors of the foot and ankle are active in 700/0 of the running cycle. The intrinsics, plantar flexors, and peroneals are important stabilizers of the plantar surface and hindfoot
during the foot nat phase (Aclelaar, 1986).
I
I
Pretibial muscle
Triceps (calf)
I
-.1.,Sw;n9....1
phase
I
1
I
phase
I
I
.......INlVNm'fM/'lir
-NjWl~----
.L--../MlWWJW'J,IIlWil,J----...L
I
1
Foot I
I
'"""-1 li Jl
Inilial
MOTION OF THE TARSAL BONES
I
1__ Slance
floor
contact
Lift
off
Initial
floor
contact
Subtalar Joint Motion
The joint between the talus and calcaneus is
lermed the subtalar joinl. Its complex motion in
three planes produces the motions of supination
and pronation, c1inicnll.y referred to as subtalar in-
Schematic phasic activity of leg and foot muscles during
normal gait.
in it study enlailing eXlkTinh..'lllal sL'kClive fusio
thl:sC joints. Sliblala,· joint arthrodesis redllcL'd ta
~l\·iclilar Illolion to 26% or its normal mol ion and
duced calcaneocuboid mOlion lO 56~~ normal. Ca
neocuboid arthrodesis n.:duccd sublala,· motio
92 % normal and talonavicular motion to 67':'/c
mal. Selective fusion of the lalomwicuhu- joint
the most profound effect on the remaining joillls
ducing their relT'laining Illotion to only 1" each. B
doin (1991) showed thal "xperimelllal slIbtala
sion resulled in Significant reduction in talonavic
joil1l contact and nlso reduced ankh.: joil1l contac
A
Tarsometatarsal and Intertarsal Motion
B
Subtalar joint axis. A, Sagittal plane (lateral view). The axis
rises up at a
4r angle from
the plantar surface. B, Trans-
verse plane (top view). The axis is oriented 16° medial to
the midline of the foot. Reprinted with permission from Manrer, J. T. (7947), Movements of the subia/af and transverse tarsal
joints. Ani!! Rec. 80, 397.
transverse tarsal joint: a longitudinal axis and an
oblique axis. The longitudinal axis is oriented lSo upward from the horizontal and go medially frolll the
longitudinal axis of the foot. Inversion and eversion
occur about thc longitudinal axis (Fig. 9-16). The
oblique axis is oriented 52° upward from the horizontal and 57(> anlcromcdially. Flexion and extension
OCClIr primarily aboLit this axis (Fig. 9-17). Ouzonian
and Shereff (1989) determined in vitro talonavicular
motion to be 7° in nexion-extension and 17u in pronation-supination. Calcaneocuboid motion is 2{' flexionextension and 7° pronation-supination.
The motions of the subtalar joint and transverse
tarsal joint interrelate to produce either foot flexibility or rigidity. Elftman (1960) showed that the
major axes of the calcaneocuboid joint and talonavicular joint are in parallel when the subtalar joint
is everted, thus allowing motion of the transverse
tarsal joint. As the subtalar joint inverts. the axes of
these joints are convergent, thus locking the transverse tarsal joint and providing rigidity to the midfoot (Fig. 9-18). During the period of midstancc to
toe-off, the foot therefore becomes a rigid lever
through inversion of the subtalar joint and locking
of the transverse tarsal joint.
Astion (J 997) showed the close interplay between
the subtalar, talonaviclllaJ~ and calcaneocuboid joints
The joillls between lhe three clIneiforrns, cub
and five melatarsals produce little motion. The
terlarsal joints are c1oscl~· congruent and exh
minimal gliding motion bct\\'cL'n OIlL: anolher.
tarsometatarsal joints, known as Lisfranc's jo
,
,~
/'
/
/.'\.
,,
:,
,,
,
Comparison of the posterior calcaneal facet of the righ
talar joint with a right-handed screw. The arrow repres
the path of the body following the screw. The horizont
plane in which motion is occurring is hh'; n' is a plane p
pendicular to the axis of the screw; s is the helix angle o
the screw, equal to s', which is obtained by dropping a
pendicular (pp') tram the axis. As the calcaneus inverts,
tates clockwise and translates forward along the axis.
Reprinted '..vith permission from M<1fller, J. T. (1941). Moveme
of thl? subta/ar and lransvers£' tarsal joints. Ana! Rec. SO, 391
Talonavicular
axis
A
9;t~'--~d
..-:---
Lat.
Navicular
Calcaneus
B
Normal
Longitudinal axis of the transverse tarsal joint. Inversion
and eversion occur about this axis. A. lateral view. B. Top
view. Reprinted wirh permission from lv/anter. J. T. (194/).
Anteroposterior view of the transverse tarsal joint of th
right foot. The anterior articulations of the talar head a
calcaneus are shown. The major axes of the talonavicul
and calcaneocuboid joints are shown in the neutral pos
tion (parallel) and with the heel in varus (convergent).
Movements of the svbtalar dnd transverse tarsal joints. Anat
Rec. 80, 397.
Q~~----------------
are intrinsically stable because of their arch-like
configuration best seen in cross-scClion. The sec-
ond metatarsal base is recessed into the mid foot,
forming a key-like configuration with the intermediate cuneiform (Fig. 9-19). A strong ligament
Varus
•
known as the Lisfranc's ligament connects the s
ond metatarsal base to the medial cuneiform. T
motion of the first three metatarsocuneifo
joints is minimal compared with the fOLlrth a
fifth melatarsocuboid joints, Ouzonian and Sh
efT determined the first metatarsal-med
cuneiform motion to be 3.5 0 f1cxion-extension a
1.5 0 pronation-supination, while the fourth a
fifth metatarsotarsocuboid joints \Vcre 9 to
flexion-extension and 9 to 11° pronation-supin
tion. An in vivo study or first mctatarsocuneifo
motion found a mean sagittal motion of 4.4(' (Fr
& Prieskorn, 1995), An in vilro sludy of fi
metatarsocuneiform motion showed that on
llOlo of lOa specimens demonstrated an,Y addu
tion-abduction (vVanivenhaus & Prcttcrklieb
A
1989), There has been recent conccrn that hyp
mobility of the first metatarsocuneiform jo
may lead to offloading of the first ray and sub
bar.
!
-------------------
Oblique axis of the transverse tarsal joint. Flexion and ex~
tension occur about this axis. A. lateral view. B, Top view.
Reprinted wirh permission from Manter. 1. T. (1941). Movemenrs
of the subta/r1r and transverse tarsal foints, Anat Ree. 80, 397.
quent hallux valgus deformity (Klaue, Hans
& l\IIasquelel, 1994), Mizcl (1993) described
plantar first metatarsocuneiform ligament as
major restraint to dorsal angulation ancl sllb
quent dorsal displacement of the distal fi
metatarsal head.
Motion of the Hallux
The hallux must accommodate a \vide range of m
lion of the foot to perform a great variety of tas
Dorsal
Phalanx
Lateral
O'
40
70~
Plantar
Med
Llslranc's JOInt
Dorsal
Melatarsal
Medial
B
A
Plantar
A, Top view of the tarsometatarsal joints, known as Lis·
franc's joint. Note the recessed second metatarsal base. B,
Cross-sectional view of lisfranc's joint seen on computed
tomography scan. Note the arch-like structure.
!
!
Late
Contact distribution of the first metatarsophalangeal
in 0<> (neutral), 40° of extension. and 70<> of extension
Joint contact of the proximal phalanx. Bottom, Joint
tact of the metatarsal head. With increasing extensio
1
.__i_o_in_'_'_U_rf_a_,_e_,_o_n_,a_'_ls_'h_i_f'_d_o_,_,a_1_IY_O_n_'_h_e_m_e_,_a_'d_,_,a_l
Normal Weight Bearing
Normal Weight Bearing
7
Medial
Lateral
collateral J;:~",,~'--1~rcollateral
ligament -
ligament
Instant
center
A
B
A, Instant center and surface motion analysis of the
metatarsophalangeal joint of the hallux in the sagittal
plane. Each arrow denoting direction of displacement of
the contact points corresponds to the similarly numbered
instant center. Gliding takes place throughout most of the
motion except at the limit of extension, which occurs at
toe·off in the gait cycle and with squatting. At full extension, joint compression takes place. The range of motion
of the hallux is indicated by the arc. B, Instant center
analysis of the metatarsophalangeal joint of the hallux in
the transverse plane during normal weight·bearing. Gliding (denoted by arrows) occurs at the joint surface even
though the range of motion is small.
One nced onl.\' consider the dorsif1exed positio
the gr.:at toe or n baseball catcher crollching be
home plate to appreciate the degree of motion p
ble in the great lOe. The first metatarsophalan
joint has a range of motion from 30" plantarllcxi
90" dorsiflexion with rL'slk'Cl to the long axis o
first metatarsal shaft. The first metatarsal is inc
20" with respect to the noor; thercfore, range of
lion or the hallux is 50" plantarllc:don to 70" d
llexion referenced to the lloor surface. During
(oc·off phase of nanna I walking, maximum dors
ion of the nrst l1lt.:tatarsophalangeal joint is requ
Anal~'sis of motion of the hallux in the sa
plane reveals that inslallt centers of motioll
fall within the center
the metatarsal head,
minimal scatter (Fig. 9-20). The surface motio
the first metatarsophahll1gcal joint is chan\clC'
as tangential sliding from maximum plantarlle
to IT'lOdcrate dorsiflexion. with some joint com
sion <:\t maximurn dorsillexion (Sammarco,
Shc,dL Bcjahi, & Kummc,; 1986). Aim el aL (1
determined the first metatarsal head surface co
area to be 0.38 l'm~ in the neutral position, w
dec"eases to 0.04 em' in full dorsiflexion (Fig. 9
The metatarsal head surface contact area shi fts
sail\' with full c!xtension and is associated with
con~prcssion. This explains the characteristic
nwtion of dorsal osteophytcs and limited dors
ion or the proximal phalanx in cases or h
rigidus (Fig. 9-12).
or
A
50-year-old woman who
has been wearing shoes
with a narrow Ioebox for <llmost
35 years. By compressing the
forefoot medially and laterally,
these abnormal forces can lead
to a hallux valgus deformrltion.
In this way, the proximal pha-
lanx shifts laterally and pronates
on the first metatarsal head.
This abnormal position of the
proximal phalanx decreases its
ability to depress the metatarsal
.
lateral view of hallux rigidus. Note osteophytes on the
dorsal aspect of the metatarsal head, which limit joint extension .
------------------
The great toe provides stability to the medial aspect of the foot through the \vindlass mechanism of
the plantar aponeurosis (sec The Medial Longitudinal Arch). As the body passes ovcr the foot in toeoff. the proximal phalanx passes over the metatarsal
head and depresses it. This has been conflrmcd by
force plate analysis in the lutc sLancc phase, which
shows thai pressure under the first metatarsal head
increases in this phase of gait (Clark, 1980). In hallux valgus, the proximal phalanx shiFts laterally ancl
pronates on the first metatarsal head. Thb abnOl-·
Illal position of the proximal phalanx decreases its
head during toe-off (Fig. 9-23).
A view of the plantar surface of the foot of a patient with
severe hallux valgus. Note calluses underlying the second
and third metatarsal heads (transfer lesions), which indicate a transfer of plantar forces away from the first
metatarsal head to the lesser metatarsal heads.
The altered joint mechanics produced by hallux valgus is
evident in analysis of instant centers of rotation, which
demonstrate joint distraction and jamming where gliding normally occurs (Fig. 9-24).
Abnormal Weighl Bearing
(Bunion)
ability to depress the metatarsal head during toeoff. Any clinical situation that affects the normal depression of the metatarsal head may t1'ansfer plantar forces laterally to the second and third
metatarsal heads and rcsulL in the formation of
painful plantar calluses called transfer lesions (Case
Study 9-1 and Figs. 9-23 and 9-24).
Motion of the Lesser Toes
The lateral four loes arc analogous to the digits of
the hane!. The lesser tocs have three phalanges. the
motion of each being controlled by extrinsic muscles. which originate within the leg, ;;lnd intrinsic
muscles, which originate within the fool. Normal
motion of the mctatm·sophalangeal joint is approximatelv 90" eXlension to 50" flexion. The extrinsic and intrinsic muscles contribute to the toc
extensor hood. which controls motion of the
metatarsophalangeal and interphalangeal joints
- ' ' ' - - - - - - - - - - - - - - - Ftoor
Altered instant centers caused by the presence of a
bunion, seen through analysis of motion of the metatarsophalangeal joint of the hallux in the sagittal plane.
.I Each arrow denoting direction of displacement of the
contact points corresponds to a similarly numbered in·
stant center. The arc indicates the range of motion of the
hallux, which is more limited than that in a normal foot.
Extensor digilorum longus relaxed
!
Extensor digitorum
longus tendon
Interosseous
muscle
I
Extensor
Sling
hood
Flexor digilorum
longus tendon
insertion site
Extensor sling
Extensor digitorum
longus contracted
Lumbrical
tendon
Deep transverse
ligament
Flexor
tendon
sheath
1
Flexor digitorum
Sling traction
brevis tendon
Lateral view of the lesser toe illustrating the extensor
Lateral diagram showing the action of the extensor s
When the extensor digitorum longus contracts (botto
the proxima! phalanx is lifted into extension through
sagittal bands.
hood with contributing muscles and ligaments.
(Fig. 9-2S). The e'trinsics consist of the long toe
nexors and extensors. The lumbricals and interossei are the main intrinsic contributors to the extensor hood. The intrinsics act to l1ex the metatarsophalangeal joints and extend the inter~
phalangeal joints (Fig. 9-26). The long toe cxten~
sors extend the metatarsophalangeal joint through
the action of the sagittal bands by lifting the pmximal phalanges into e'tension (Fig. 9-27). The
flexor digilorum brevis is rhe primary Oexor of the
proximal interphalangeal joint. The ne,or digitorum longus is the primary flexor of the distal interphalangeal joint.
Neurological conditions, such as diabetic neuropathy or Charcot-l\tlarie-Tooth disease, initially
affect the intrinsiL: muscles or the fOOl and res
an intrinsic minus condition. The extrinsiL' lllu
overpower the intrinsics and a claw !C)C dcform
produced with extension of the mt.:tatars
langeal joint and flexion of the interphala
joints (Fig. 9-28).
Interossei
Extensors
I
Extensors
-
I
Flexor
digitorum
longus
Flexor
digitorum
Lumbricals
~brev!s
Interossei
Lumbricals
The intrinsic muscles (interossei and lumbricals) act to flex
the metatarsophalangeal joint and extend the interphalangeal joints.
I
-------------
A claw toe deformity is produced by an imbalance o
extrinsic and intrinsic muscles. Relative weakness of t
terossei and lumbricals with overpull of the extrinsic
extensors and flexors produces an intrinsic minus def
mity of metatarsophalangeal extension and interpha
langeal joint flexion.
c
W
2
W
2
GE'-----------The beam model of the longitudinal arch. The arch is a
curved beam consisting of interconnecting joints and supporting plantar ligaments. Tensile forces are concentrated
on the inferior beam surface; compressive forces are generated at the superior surface.
I ---------------e
THE MEDIAL LONGITUDINAL ARCH
models e~ist to describe the medial longitudinal arch of the foot: the beam model and the truss
model (San-anan, 1987). The beam model slates
that the arch is a cun'cd beam made up of interconnecting joints whose structure is dependent on joint
and ligamentous interconnections for stability. Tensile forces are produced on the inferior surface 01"
the beam and compressive forces are concentrated
on lhe superior surface of lhe beam (Fig. 9-29). The
truss model Slales that the arch has a triangular
structure wilh two slnlls connected at the base by a
'[\1,10
tic rod. The SlI-uts arc under compression and
tie rod is under tension (Fig. 9-30). BOlh mod
h(:t\·c validity and can be demonstrated clinically.
The structure analogous to the tic rod in
truss model is the plantar fascia. The plantar fas
originates on the mediallubcrosily of the calcane
and spans the transverse tarsal, tarsometatars
and metatarsophalangeal joints to insert on
metatarsophalangeal plantar plates and collate
ligaments as well as the hallucal sesamoids. Do
nexion of the metatarsophalangeal joints pla
traction on the plantar fascia and causes elevat
of the arch through a mechanism known as
"windlass effect" (Hicks, 1954) (Fig. 9-31). Dur
toe-off in lhe gail cycle, lhe loes are dorsiflexed p
sively as the body passes over the foot and the pl
tar fascia lightens and aels to shorten the distan
between the metatarsal heads and the heel, th
elevating the arch. The traction on the plantar f
cia also assists in inverting the calcaneus lhrou
its altachment on the medial plantar aspect of
calcaneus.
The arch has both passive and active suppo
I-luang et al. (1993) performed an in vitro study
the loaded fOOl and foundlhat division of the pl
tar fascia resulted in a 25°M decrease in arch st
ness. They found the three most important sta
contributors to arch stability in order of imp
A
o
B
A, Schematic of a truss. The far left wooden segment re
resents the hind foot, the middle wooden segment repre
sents the forefoot. and the far right wooden segment is
the proximal phalanx. The rope is the plar"!tar fascia. B.
Dorsiflexion of the proximal phalanx raises the arch
through traction on the plantar fascia.
tance were the plantar fascia, the long and shon
platHar ligaments, and the spring ligament (cakaneonavicular ligament). The spring ligament forms
a sling for the talar head, which prevents medial
and plantar migration of the talar head, and therefore provides stalic arch support (Davis ct aI.,
1996). Basmajian (l963) demonstrated e1ectromyographi-cally that the muscles of the calf did not
contribute to support of the medial longitudinal
arch when a load was applied to the leg of a seated
individual; however, his experimental model did
not simulate normal walking or running, when
arch integrity is under grealCr challenge. Thord~l.I·­
son et al. (1995) performed a dynamic study of arch
support by simulating stance phase of gait by applying proportional loads 10 tendons while the foot
was loaded. [n this study, the plantar fascia contributed the most to arch stability through toe dorsinexion. The posterior tibialis contributed lhe
most lO dynamic arch SUpPOrl. In a similar cadav~
eric study, Kitaoka ot al. (1997) demonstrated a
0.5 mm decrease in arch height and an angular
change in lhe bones of the arch when tension on
the posterior tibial tendon was released during simulated stance. A clinical study of 14 feet postplantar
fasciotomy at greater than 4 years follow-up
showed a decrease in arch height of 4.1 111m, thus
supporting the truss model of arch stability (Daly et
aI., 1992).
Neuropathic joint changes or trauma may disrupt
the bone and joint support of the arch, leading to
. arch collapse and a resultant rocker bottom foot deformity (Fig. 9-32). This sequela of joint destruction
lends credence to the beam model of arch stabilitv.
A lateral radiograph of a rocker bottom foot, dem
ing loss of the bony and Iigamento\ls support of th
.-._._----_._-----_._--
or
bod,v weight forward ~l1ullg the il:ods
prC)f
(F;ig. 9-33). Thc mOll1cnt produced by cach
tendon unit call be pn:dicted b,\' their rclati
to the ankle ~HH.l subtahir a,'\(.'s (Fig. 9-34).
The soleus and gaslnK!lr.:-l1lius combine t
the ,Achilles tendon, which insens onto the
neus and is the strongest fk',,\or of the ankle. ;
ematical Inodel has prcdicted peak Achilles
forces 10 be 5.3 to 10 til1lt.·s bod.\' wt.·iglll durin
ning (Burdett, 1982). Firing or the ankk' plan
aI's during midstancc acts to slow tht.' fon\'a
lion of the tibia on:r lIw fool.
MUSCLE CONTROL OF THE FOOT
Twelve of the thirteen extrinsic and nineteen intrinsic muscles control the foot and ankle. The plantaris
muscle is an extrinsic muscle that generally has no
contribution to muscle control or the foot or ankle.
The extrinsic muscles are the strongest and most
important in providing active control during gait.
According to Fick's principle (1911), the strength of
a muscle is proponional to its cross-sectional area.
Accordingly, Silver et al. (1985) have weighed and
measured muscle fiber length to determine the relative strengths of muscles acting on the foot and ankle (Table 9-1).
The muscles of the leg fire in a pattern during
normal gait to ensure an efficient lransrcr of mw;de force to the floor and smoolh progression of"
The Relative Strength5 of MU5Cles Acti
on the Foot and Ankle
Plilntarflexor
Strength Percentage
Dorsiflexor
Strength Percentage
Soleus 29.9
TibIaliS anterior 5.6
Gastrocnemius 19.2
Extensor digitorum long
Flexor halluc is longus 3.6
Extensor hallucis longus
Flexor digitorum longus 1.8
Peronel.Js tertius 0.9
Inverters
Everters
Tibialis posterior 6.4
Peroneus longus 5.5
Peroneus brevis 2.6
------------
Walking Gait
8lance phase
Swing phase
~---------I~--~---'
Muscle
Heel
slrike
Foot
flat
Heel
all
Toe off
I
I
I
Tibialis anletior
Ext. digitorum longus
Ex!. hallux longus
Gastrocnemius
Tibialis posterior Flexor digitorum longus
Flexor hallux longus
Peroneus longus
Peroneus brevis
Abductor hallux
Flex. hallux brevis
Flex. digilorum brevis
Abd. digit minimi
Interossei
Ext. digitorum brevis
I
I
I
I
I
I
I
I
I
I
I
I
i
I
I
I
I
I
I
I
I
I
!
!
!
I
,
,
10% 20% 30% 40%
50~;'
I
60%
,
70~o
80~o
90% 100%
Electromyography ollhe Foot During Walking
Electromyography of the musculature of the foot and ankle during one normal gait cycle
(heel strike to heel strike) .
•
The strongest extensor of the ankle is lhe tibialis
anterior, which is 1110st active during stance phase
from heel strike 10 foot nat. The ankle and toe extensors fire eccentrically to slow the descent of the
foot and prevent foot slap. They also are necessary
to allow fOOl clearance fTom lhe noor during the
swi ng phase.
The strongest inverter of the foot ancl ankle is the
posterior tibialis muscle. The postcrior tibialis is a
dynamic supporter of the medial longitudinal arch.
Il flmctions to invert the subtalar joint during midand latc stance, thereby locking the tranSVel'se tal'Sal
joint and ensuring rigidity of the foot during toe-all
Loss of this muscle results in acquit'cd pes planus
with Ilattcning of the arch, abduction of thc forefoot.
and evcrsion of thc heel (Fig. 9-35). Patients with
posterior tibialis tcndon dysfunction usually are unable to activcly invert their heel while attcn1pting a
single toe rise. They have difficulty performing a sin·
gle toe rise because of their inability to form a rigid
platform on which to support their weight.
The primary evertcrs of thc foot and ankle arc th
peroneals. The peroneus longus inserts on the ba
of the first metatarsal and Illcdial cuneiform an
acls to depress the metatarsal head. Injury or para
ysis of this muscle may allow elevation of the fir
metatarsal head and decrease loads borne by th
first metatarsals and can rcsult in the developme
of a dorsal bunion. Thc peroneus brevis stabiliz
the forefoot laterally by resisting inversion and w
found by Hinlermann and associates (1994) 10 b
the strongest everter of the foot. Loss or perone
muscle strength can result in varus of the hindfo
(Sammarco, 1995).
The interosseus muscles are active during la
stance and are thought to aid in stabilizing the for
foot during toe-olT. An imbalance between the i
Irinsics and extrinsics will lead to toe deformiti
such as hammer toes, claw toes, or mallei toes.
,Both intrinsic and extrinsic muscles; mediate th
positional control of the great toe. A cross·section
the proximal phalanx shows the relative position
Longitudinal
axis
Sublalar
axis
I
• I.
I
I
EHL
I
I EDL
I
I
I
I
I
I
Ankle axis
TP.
.
FDL.
Loss of the medial longitudinal arch in an acquired a
flatfoot secondary to posterior tibial tendon deficien
'------------------
Subtalar and ankle axes in relationship to extrinsic muscles.
EDL, extensor digitorum longus; fHL, extensor hallucis
longus; FDL, flexor digitorum longus; FHL, flexor hallucis
longus; PS, peroneus brevis; PL, peroneus longus; TA, tib·
ialis anterior; Te, tendon calcaneus; TP, tibialis posterior.
man absorbs 63.5 tons on each foot while wa
Running one mile \\'ould produce 110 tons per
that same 150-lb man (Mann, 1982).
ivlanter measured the cOr'npl'cssive loads
static loading in cadaveric feet lo determine th
lribution or forces through the joinls of th
(Fig. 9-37). The highest pan or the longit
arch, the talonavicular and naviculocune
Normal Tendon Position
.......... Extensor hallucis long
the nexors, extensors, abduclors, and adductors
(Fig. 9-36). The tibial and fibular sesamoids lie
within the toe tendons of the flexor hallucis brevis
muscle, beneath the head of the first metatarsal.
Similar to the patella, they increase the lever arm
distance of the pull of the newr halluc is brevis
muscle and enable greater flexion torque to be generated at the metatarsophalangeal joint. They also
act to transfer loads from the ground to the first
metatarsal head.
Kin.etics
of the
Foot
The magnitude of loads experienced by the foot is astounding. Peak vertical forces reach 120°10 body
weight during walking, and they approach 275°10 during running. It is estimated that an average 150-lb
?-'""-"'--..
Extensor hallucis br
Lateral
Medial
Abductor
hatlucis
Adductor
hallucis
Flexor hallucis ----{)-r~-.;:_cJ_-_
brevis
Flexor hallu
brevis
I
I
l
Flexor hallucis lonQus
~matiC
cross·section of the proximal phalanx
hallux showing normal positions of the various tend
relation to the bone.
The distribution of loads under the foot durin
stance has been the subject of intense investigatio
for the last half centul)'. Ini1iall:\', the concept of
"transverse metatarsal arch" was promoted, i
which loads were borne primarily by the heel, firs
and fifth metatarsal, as if the foot were a tripod
This concept was disputed by Morton (1935), wh
thought the forefoot had six contact points tha
shared equally in weight distribution, namely, th
two sesamoids and the four lesser metatarsal heads
Recent plantar pressure studies b.y Cavanaugh et a
(1987) of subjects standing barefoot have deter
mined that the distribution of load in the foot is a
follows: heel 6W/o, midfoot 8(Ye, forefoot 28%-, an
toes 4% (Fig. 9-38). Peak pressures under the hee
are 2.6 times greater than forefoot pressures (Fig
9-39). Forefoot peak pressures occur under the sec
ond rnetatarsal head (Fig. 9-40).
Static foot radiographic measurements fail to pre
dict 65(Jr) of the variance found among d~'namic pres
sures measured in various subjects. Therefore, th
dynanlics of gait exert the primary' inOuence o
plantar pressure during walking (Cavanagh et a!
1997). Hutton et al. (1973) studied the progression o
the center of pressure across the sole of the foot du
ing gait (Fig. 9-41). During barefoot \valking, th
center of pressure is initially! located in the centra
heel and accelerates rapidly across the midfoot t
reach the forefoot, where the velocity decrease
Peak forefoot pressures are reached at 80% stanc
phase and are centered uncleI' the second metatarsa
15
25
40
50
•
-'---------4.5
i
Compressive forces of the foot after a 60-1b load is applied
to the talus. The majority of the force passes through the
I talonavicular joint and into the first through third
~ __m_e_t_at_a_rs_a_'s_.
_
joints bear the majority of the load through the
tarsal joints. The medial column of the foot, consisting of the talus, navicular, cunei forms, ancl first
through third metatarsals, bears the majority' of the
load. The lateral column, made up of the calcaneocuboid joint and lateral two metatarsals, transmits the lesser load.
I
Mean regional weight distribution expressed as a percen
age of total load carried by the foot in barefoot standing
Over 60% of the weight is distributed in th~ rearfoot, 8%
in the midfoot, and 28% in the forefoot. The toes have l
tle involvement in the weight-bearing process.
132.6
At toc·oIT, the center 01" pressure is located unde
hallux. The metatarsal heads arc in contact wit
1100r at least 50% of Slance phase. Soames (1985
termined that the highest peak pressure and gre
foot-floor impulse during barefoot walking wa
del' the third metatarsal head insLCad of the sec
The distribution of plantar pressures changes
138.9
Lateral
53.4
£
t:
shoewcar. Shocwcar reduces peak heel pressu
producing a 1110re even distribution of pressure u
the heel. With shoes, forefoot load distribution
medially with maximum pressure under the firs
second metatarsal heads. The pressures unde
lOes also increase with shocwear (SoalTIes, 1985
The distribution of plantar pressure during
ning has identified two t~!PCS of runners chara
ized by their first point of contact with the gro
rcarfoot strikers and mid foot strikers (Fig. 9
Rcm-root strikers make initial ground contact
the posterior third of the shoe. The initial conta
the midroot strikers is in the middle third o
shoe. In both groups, first contact occurs along
later,al border of the foot. Peak pressure does no
fer between runner types. The cellter of pressu
in the dista!·most 20 to 40°;(; of the shoc in both
tact groups for I110St of contact timc, indicating
time is spent on the forefoot (Canuulgh ('t al.. 1
27.8
51.8
.
,.
Mean regional peak pressures during standing measured
in kilopascals (kPa). The ratio of peak rearfcat to peak
forefoot pressures is approximately 2.6:1.
80
x·
Med
70
60
"
50
5
40
<l.
~
~
"'"'
Ii:
x
~
Lat
30
20
10
0
1000;/" M
0% Lal
Foot Width
B
A
Metatarsal head pressure distribution during standing. A. A
line (XX') drawn in the contour plot between the approximate locations of the first and fifth metatarsal heads. B.
The distribution of pressure along the metatarsal head
(XX') indicating maximum pressure under the second
metatarsal head.
",'jiI.
l
,!t
.. 55%
\50%
45o~,\40%
35'0 ,30%
,,
,,
25%
,
t,
,
,:
20%
15%
+,
,,
t 10%
:
r, 5"
,
10
;2%
I
1mIIIIL-.I
_
The progression of the center of pressure along the sale of the
foot during normal walking is expressed as a broken line. Each
point on the sale corresponds to a percentage of the gait cycle. Note the rapid progression across the heel and midfoot to
reach the forefoot. where most of stance phase is spent. It
then progresses rapidly along the plantar aspect of the hallux.
During walking and running. several forces arc acting between the foot and the ground: vertical force.
fore and aft shear (anteroposterior shear), 111edial and
lateral shem; and rotational torque (Fig. 9-43). The
vertical ground reaction force exhibits a double peak
following the initial heel strike spike, The first peak
follows heel strike in early stance and the second peak
occurs in IDle stance prior lo toe-olT. The fore and aft
shear forces dcmonstt'atc initial braking by the foot as
the foot places a fonvard shear force on the ground.
followed by a backward shear on the ground as it
pushes off in late Slance. Most of the medial-lateral
shear is directed laterally because the body's center of
gravity' is oriented medially over the rool. Medial (internal rotation) torque is generated early in stance as
the tibia internally rotates and the root pronates, followed by laleral (external rotation) torque as the leg
externally rotates and the foot supinates.
lached as is evident by the somctirncs dramatic do
sal foot swelling found during trauma or infection o
the foot or ankle. The plantar skin is firmly attache
to the underlying bones, joints, and tendon sheaths o
the heel and forefoot by specialized extensions of th
plantar fascia, This function of the plantar fascia is e
sential for traction between the floor and the foot
weight-bearing skeletal structures to occur. Durin
extension of the metatarsophalangeal joints. thes
plantar fascial ligaments restl;ct the movement o
skin of the forefoot and pia mar metatarsal fat pa
(Bojsen-Moller & Lamoreux, 1979).
The heel pad is a highly specialized struclure de
signed to absorb shock. The average heel pad area
23 cm~, For the average 70-kg man. the heel loadin
pressure is 3.3 kg/cm~. which increases to 6 kg/em
with running, At a repetition rate of 1,160 heel im
pacts per mile. the cumulative effect of running
impressive. These cumulative forces would no
mally result in tissue necrosis in other pans or th
body (Perry, 1983). The heel pad consists of comma
shaped or U-shaped fat-filled columns arrayed ven
cally. The septae arc reinforced internally with cla
Mid
Rear
Fore
A
Midfoot strikers
n=5
o
(~r:J(_~~)t
x- -.
x
'------1...:..:..
Rear
-i.,-"
)
Mid
Fore
B
SOFT TISSUES OF THE FOOT
The soft tissues of the foot arc modified to provide
traction. cushioning. and protection to the underlying
structures. The dorsal skin of the foot is loosely at-
Two types of runners characterized by initial ground contact. A, Rearfoot strikers. H, Midfoot strikers.
Percenlage of Cycle
Force
(0, ot bOdy
wDi('hn
o
15
30
60
HS
FF
HO
TO
I
100
60
Vertical
-
-/" I'\.
\
20
1\0.
0 -Fore
t
Shear
AFT
Medial
20
t
10
j
Lateral
Torque (Nm)
t
j
Lateral
0
j
Stlear
Medial
20
0
,- ....
IA
.... .....
10
HS
,
\
......... V
I
I
FF
HO
I
i
TO
9
0
9
Ground reactive forces acting on the foot during the gait cycle. HS, heel strike; FF, foot
flat; HO, heel-off; TO, toe-off. Reprinted ~vitfJ permission from Mann, R.A (7982). Biomechanics of running. In AAOS Symposium on the Foot and Leg in Running Sports (.0.0.30-44). St. Louis:
C V Mosby Co.
•
tic transverse and diagonal fibers to produce a spiral honeycomb effect (Fig. 9-44). The multiple small
closed cells arc arranged to most effectively absorb
and dissipate force. \,Vith age, septal degeneration
and fat atrophy occur, which predispose the calcaneus and foot to injury (Jahss et a!., 1992a,b).
Ankle Joint Biomechanics
KINEMATICS
The ankle mortise forms a simple hinge consisting
of the talus, medial malleolus, tibial pia fond, and
lateral malleolus. The talus is shaped like a truncated cone, or frustunl, with the apex directed medially (tnman, 1976). The talus is 4.2 mm wider
anteriorly than posteriorly (Sarrafian, 1993a
single ankle joi.llt axis has been described as
ing just distal to the mcdi,il malleolus and jus
tal and anterior to the lateral malleolus (In
1976). This empirical "general" ankle axis ca
estimated by palpating the tips of the ma
(Fig. 9-45). The single ankle axis is angulated
terolaterall y' in the transverse plane and infe
erally in the coronal plane. Several authors
disputed the theory of a single axis of ankle
tion and have described multiple axes of moti
the ankle moves from dorsiflexion to planta
ion (Barnett & Napier, 1952; Hicks, 1953; H
mann & Nigg, 1995; Lundberg et aI., 198
Sammarco et a!., 1973). Barnett and Napier (
describe a dorsiflexion axis inclined down
Structure of a normal heel pad as seen on magnetic resonance imaging (MRI), A, Lateral
view. Note vertically oriented fat-filled columns. B, Top view of the heel pad demonstrating the spiral structure of the septae, which separate the fat-filled cells .
•
and laterall.\' and a plantarflexion axis angled
downward and medially (Fig. 9-46). The ankle
joint axes for dorsillexion and plantarflcxion cliffeI'
by' 20 to 3D" in the coronal plane but remain parallel in the transverse plane.
A small amount of talnI' rotation occurs during
ankle motion, which varies with axial load. Lundberg
X'
X
~
,~
Y-~Y
~Z'
Iam
_
The empirical axis of the ankle joint estimated through
palpation of the malleoli. The axis angles downward and
posteriorly, moving from medial to lateral.
Ankle joint axis variation. Top, In dorsiflexion (DF) the a
of motion XX' is inclined downward and I~terally. Midd
In neutral the axis of motion YY' is almost horizontal. B
tom, In plantarflexion the axis ZZ' is inclined downward
and medially.
el al. (I 989a-<l) used slereophologrammell)' 10 measure talar rotation during motion of the weightbearing ankle in normal volunteers. The talus externally rotated 9° from neutral to 30° dorsiflexion.
From 0 to 10" plantarflexion, the talus internally 1'0tatcd 1.4", followed by external rotation of 0.61> at
30" planlarllexion (Fig. 9-47). An in vilro slud~' of
loaded ankles demonstrated 2.5<" external rotation in
25(0 dorsiflexion, and < I" internal rotation at 3Y'
plantarllexion (Michelson & Helgemo, 1995.)
c
.2
x
"
S!
20
0
10
~
;;;
"9
0:;
0:;
"
"
c
0
·x
~
~
10
B
6
4
ca> c
2
0
a: -
-2
"0
-4
-6
"iii
-B
E
-10
w
rl--~i-~--~-~I--~i
-~i
-12
PF 30
20
10
0
10
20
30 DF
x"
Input foot position
Horizontal rotation of the talus around the vertical axis at
different positions of ankle dorsiflexion·plantarfJexion.
Moving from plantarflexion to dorsiflexion, the talus ini·
tially internally rotates slightly, then externally rotates
markedly.
&.$
I
10
20
60
Percentage of Cycle
100
100
_
Range of ankle joint motion in the sagittal plane durin
level walking in one gait cycle. The shaded area indica
variation among 60 subjects (age 20 to 65 years). Rr:pr
'..viti! permiSSion from Stiluffer. P. N.. Chao.
E- YS., S Brew5:e
norma!. d,seas
RL {1971i. Force (Jocl molion analySIS 01 the
I
•
c1fl(/
pro5the;;c anNe JOIn!s Ci:~l Onhcp. 127. 189
ankle planlarJlc.'\ion is actual!." occurring dista
the ankle itself. 'rhis midfoot motion e.'\plains
apparent ability
the foot to dorsillcx ~llld p
tarnex following ankle fusion. It also explains
abilil~'
dancers and gymnasts to align the
with the long a:"is of the leg during toe point. S
marco and associates (1973) found ave
non-weight-bearing ankle mol ion measured r
ographically 10 be 2-1." dorsiOexiol1 and 24" p
tarlk·xioll.
The normal pallcrn or ankle motion has b
studied extensively (Lalnon'aux, 1971; Murra
aI., 1964; Staullel· et aI., 1977; Wright et aI., 19
At hed strike, the ankle is in sligh I plantarflc:
Plantadlexion increases until foot llat. bu( (he
tion rapidl~' reverses to dorsilk:xion during
stance as the bod.\' p,ISSCS O\'el' the root. The mo
thel1 returns to plant<:lrllexion at we-ofr. TIl(' a
again dorsiflexcs in lhe middle or swing phase
changl,.·s to slighl plalltarflexion at heel strike (F
9-9 and 9-48). Ankle molion during normal wal
;:lverages 10.2" dorsiflexion and 14.2" plantarflex
with a total motion or 25". Maximum dorsifle
occurs at 70(;'1(.. stance phase and nwximlllll p
larJlexion occurs at loe-ofr (Slaufrer el aI., 1977
or
or
12
~.~
"
0
f-
rDm'--
siflexion (extension). A wide range of normal mo-
lion for Ihe ankle has been reported and depends
on whether the Illotion is measured clinically with
a goniOll1cter or whether it is measured radiographically. Goniomctric mCasurements yield a
normal motion of 10 10 20" dorsiflexion and 40 10
55" plantarflexion. Lundberg el al. (1989a-d) found
thal the joints of the midfoot contribute 10 to 41%
of clinical plantarflexion from neutral to 30" planlarflcxioIl. Therefore. whal appears to be clinical
0:;
0
a":
Ankle mol ion occurs primarily in the sagittal plane
and is described as planlarflexion (flexion) and dor-
~
;;;
'0
6
I
I
C
RANGE OF MOTION
"
~
·0
SURFACE JOINT MOTION
Sammarco el al. (1973) pcrronncd analyses o
slant centers or rOlation and surface velocitie
both normal and diseased ankles. They round
,,~
-0...
the instant centers of rouuion fell within the tali of
normal ankles but that their positions changed
,,'ith ankle motion (Fig, 9-49), This confirms that
the ankle axis of rotation does not remain constant
with motion, Surface motion from full plantarllexion to full dorsiflexion was also determined. Beginning in full plantarncx-ion. the ankle joint
;howecl early distraction as dorsillexion began.
Joint gliding then took place until full dorsillexion
waS reached and jamming of the joint occurred. It
is possible that distraction and jamming of the
.Otihiotalar joint playa role in lubrication
the
joint. In arthritic ankles the direction of displace-
or
,
,
,
\
\
\
I
I
\
I
"
1
f
f
,I
\
'
I
,I
,,"
/
,
'
, I,
I
I'
I
I
,<
"
,'
\
I
I
1
"
J
5 "
_.....--1-......
"
,' _4
I
I
Anterior
f
\
I
\ 1 I
- - _ .J
Posterior
'-
ment of the conLact points showed no consistent
palLcrn. The tibiotalar joint sllrfaces distracted in
an unpredictable manner, and they jammed when
the joint was in neutral position rather than at the
end of dorsillexioll (Fig, 9-50),
ANKLE JOINT STABILITY
Stability or the talocrural JOIllt depends Oil both
joint congruency and supporting ligamentous structures. The lateral ankle ligaments responsible for resistance to inversion and internal rotation are the
anterior taloflblilar ligament, the calcancoflblilar
,,
,,
,,
,,
,
,,
,,
,,
,,
,,
Tibia
,,
AnI.
Post.
Talus
Instant center pathway for surface joint motion at the
tibiotalar joint in a normal ankle from full plantarflexion
to full dorsiflexion. All instant centers fall within the talus.
The direction of displacement of the contact points shows
distraction of the joint surfaces at the beginning of motion
(points 1 and 2) and gliding thereafter (points 3 and 4).
1
Reprinred with permission from Sammarco, G.J., Burnstein,
AH., & Frankel, VH. (1973). Biomechanics of rhe anklE': A kinemarie sWdy. OrttlOP Clin North Am, 4, 75.
Instant center and surface velocity analysis in an arth
ankle. The instant centers vary considerably. Joint co
pression occurs early in motion and distraction occurs
dorsiflexion (velocity 4).
ligamenL. and the posterior talofibular ligam
(Fig, 9-51), The superfiCial and deep deltoid
ments arc responsible for resistance to eversion
external rotation stress. The ligaments respons
for maintaining stability between the distal fi
and tibia are the syndesmotic ligaments. The
desmotic ligaments consist of the anterior tibio
lar ligarnent, the posterior tibiofibular ligament
transverse tibiofibular ligament (also referred
a deep portion of the posterior tibiofibular), and
interosseous ligament (Fig. 9·52).
The lateral ankle ligaments are the most c
monly injured and thercroh.~ the 1110st frcquc
studied. The anterior t41lofibular and calcancofib
ligaments form a I OSlO angle with one another
9-53). They act synergistically to resist ankle in
sion forces. The anterior talofibular ligament is
del- greatest tension in plantarnexion and [he c
neofibular ligament is under greatestlcnsion in a
dorsillexion (Cawley & France, 199\; Inman, \
Nigg et aI., \990; Renstrom et aL,1988), The anle
taloribular ligament therefore resists ankle inver
in plantarOexion and the calcancofibular ligamen
sists ankle inversion during ankle dorsiflexion.
accessory functions of the anterior talofibular
ment are resistance to anterior talar displacem
from the.: m()J"tisc, c1inicall.\' refclTt:d lO as anlcrio
drawel: and resistance LO int<:rmtl rotation of (h
talus within the mortise (Fig. 9-54). The calcnn
ofibular ligament spnns both the lateral ankle joi
and lateral subtalar joint. thus conlributing lo sltt)l
Jar joinl stabilitv (Stephens & S'lllll11arCO, 1992). Th
posterior talo(lbular ligamenl is under greatest strai
in ankle dorsiflexion and aCls lo limit posterior (al
displacement within the monise.: as \\'ell as limit tal
eXlernal rotation (Fig. 9-55) (Sarrafian, 1993a). I
vitro tesling of unloaded ankles subjected to alllcrio
drawer testing del1lonstrated that the anterio
talofibular ligamcill was most important in pla
tarnexion and that the calcancofibular and postel"i
taloflbular ligaments were most important in ank
dorsiflexion (l3l1ll1clI el aI., 1991 l.
Groove lor
,,:-'- Tibialis Posterior
- Tendo calcaneus
_ Groove for
FI. Halfucis Longus
Bursa
n/f.J
Navicular bone j 1
Tibia-navicular fibres·
)
Plantar calcanea-navicular .I
Jig. and associated fibres
Calcaneus
I LMediallubercle of talus
l Posterior tibio:,alar fibres
Sustentaculum tall
Tibia-calcanean fibres
Top, Lateral side of the foot and ankle. Bottom, Medial side
of the foot and ankle_ From Anderson, 1. (Ed.) (1978). Gram's
Atlas of Anatomy. Baltimore, MD. Lippincott, Williams & Wilkins.
20
D
Posterior View
Anterior View
"5
ro
"E
~
~
c
"E
'u
"c.
Interosseous
membrane
10
~
"0
Anterior tibiofibular
ligament
ci
z
0
%I!P--\t-lnterosseous
ligament
80
Posterior
tibiofibular
ligament
Components of the ankle syndesmosis.
£
I
120
140
Degrees
Inferior transverse
(tibiofibular) ligament
as
I
100
Average angle between the cafcaneofibular and talofibu
lar ligaments in the sagittal plane. The average angle is
105° with considerable variation from 70 to 1400 among
measured subjects.
<
Sprain Injury
o
ATF
A
basketball player with an injury that results from a
filII on a plantarflexed and inverted ankle position
during a game (Case Study fig. 9·2·1).
IR
B
func.tion of the anterior talofibular ligament (AIF). A. The ATF lim·
its anterior shih (AS) of the
talu~ or
posterior shift (PS) of the tibia-
fibula. B. The ATF limits internal rota lion OR} of the talus or external rotation (ER) of the fibula. PTF. posterior talofibular ligament;
0, deUoid ligament.
Clinically, lite most commonly sprained ankle ligament is the anterior talofibular ligament, followed
by the calcancofibular ligament. These injuries most
commonly occur as a result of landing or falling on
a plantarflcxcd and invcl·ted ankle (Case Study 9-2).
During periods of ankle unloading, the ankle rests in
, position of planlarflexion and inversion. rr the
ground is met unexpectedly, lateral ligament injury
occurs. Ataarian ct al. (1985) tested the strength
of the ankle ligaments by loading cadaver ligaments
to failure and found the strength of the various
ligaments from weakest to strongest to be anterior
., Case
,.. Study Figure 9-2-1.
An abnormally high road in conjunction with the loading
:: rate produces the injury (Case Study fig. 9·2-2). The sprain
inversion injury produced by high stress (load per unit of
area) in the plantarflexion and inversion direction will most
commonly aff~ the antero talofibular ligament (failure
PTF
load -139N). This produces lateral instability in ttle ankle
joint an abnormal anterior talar displacement from the mortise. decreasing the resistance to internal rotation of the talus
within the mortise.
Abrupt rupture
Microtears [
139 N
..
A
I
sprain InJury
B
~-----­
Function of the posterior talofibular ligament (PH). A, The
PTF limits posterior shift (PS) of the talus or anterior shift
(AS) of the tibia·fibula. B, The PTF limits external rotation
(ER) of the talus or internal rotation (lR) of the fibula, ATF,
anterior talofibular ligament; D, deltoid ligament.
f
k.."".......
,_..~ :00"""
.j.
associa~e~ ~vilh
"",.. ",,,,L _I!!""'l'-""""
-~
-...:.<..
Elongation of the Antero Talofibular Ligament
(:ase Study Figure 9-2-2.
_~...~~=. . . . . . . . . . . . . . ------,---~.~.~'-7.".7.=.'
..._ ...
,:,,:,.~.4,~,~~.
TI14'"
(Medial talar lill)
LS 1.9 mm
AS 5.6 mm
IT 0
....
Transection of both superficial and deep components of
the deltoid ligament resulted in an average 14 0 of valgus
talar tilt (TT) among 24 cadaver specimens.
lalofibular 139 N, posterior talofibular 261 N, calca,
neofibular 346 N, and deltoid 714 N. Therefore, incidence of ankle ligamentous injury tends to match
both mechanism of injury and ligamentous strength.
The deltoid ligament acts to resist eversion. external rotation, and plantarOexion of the ankle joint
(Fig. 9-56) (Harper, 1987; Kjaersgaard-Andersen et
aI., 1989; Nigg et aI., 1990). It also resists lateral talar shift within the mortise when the mortise is
widened by distal syndesmotic ligamentous injUl)1
or distal fibular fracture (Fig. 9-57) (Michelson,
Clark, & Jinnah, 1990; Harper, 1987).
Under physiological load. the ankle articular surface congruency takes on more imporlance (Cawley
& France, 1991; Stiehl et aI., 1993; Stormont et aI.,
1985). Stormont et al. found that in a loaded slate
the ankle articular surfaces provided 30% of rotational stability and 100% of resistance to inversionl
eversion. They hypothesized that during weightbeming. the ankle ligaments do not contribute to an·
kle versional stability. although rotational instability
may still OCClll: Cawley and France showed that the
force to cause ankle inversion and eversion increased
by 91 % and 80%, respectively, with loading. Stiehl et
al. (1993) found lhal loading of Ihe ankle resulled in
decreased range of motion (especially plantarncx.ion),
decreased anteroposterior drawer. as well as increased stability against version and rotation. Cass
LS 1.9 mm
AS 5.6 mm
IT 0
LS 3.8 mm
AS 8mm
IT 0
The lateral malleolus is excised to simulate a fibular
ture or distal syndesmotic ligament injury. A lateral
rected load is applied to the talus in an unloaded a
modeL The talar lateral shift (LS). anterior shih (AS)
valgus talar tilt (TI) are measured. S~etioning of the
deltoid (DO) doubled the lateral talar shift from 1.9
mm. SD, superficial deltoid.
•
and Settles (1994) performed CT scans of loaded cadaveric ankles in an apparatlls that did not constrain
rotation and demonstrated that talar lilt or an average
20 still occllITcd in loaded ankles aftcr sectioning
both anterior talofibular and calcaneonbular ligaments. They did not feel thal the articular surfaces
prevented inversion instability during ankle loading.
Most studies agree that loading of the ankle results in
increased stability as a result of articular surface con·
grueney. especially in ankle dorsiflexion.
Syndesmotic stability is dependent on the integrity or both malleoli. the syndesmotic ligaments,
and the deltoid ligamentous complex. During ankle
dorsinexion. there is approximately I rnm of mortise widening and 2" of external rotation of the
fibula (Close, 1956). The normal distal fibular migration with loading is I rnm (\'\fang et aI., t 996).
This distal fibular mignllion serves to deepen the
ankle mortise for addcd bony stability (Scranton,
McMaster, & Kelly, 1976). With disruption of the
mortisc in an external rotation injury, the s~'n­
desmotic ligaments and deltoid ligaments are tOI'n,
the distal fibula fmclures, and the talus displaces
laterally. A study of cadaveric ankles by OlgivieHarris et al. (1994) deflI1ed the conlribution to resistance to lateral talar displacenlcnt by the syndesmotic ligaments to be 350/0 for the anterior
tibiofibular ligament, 40% for the postcrior tibiofibular ligan1cnt, 22% for the interosseous ligarncnt, anel
less than t 00/0 for the interosscous membrane.
The deltoid ligament appears to be key in preventing lateml talar shift. Burns et al. (1993) found
only minimaltalar shift in a loaded cadaveric ankle
study with sectioning or the syndesmotic ligaments
until the deltoid ligament was sectioned. Michelson
and colleagues (1990) simulated 4 mm of lateral
fibular displacement in a cadaver study and placed
the ankle under a 100-lb load. Lateral lalar shift
doubled from I to 2 mm following sectioning of the
deltoid ligament. Pereira et al. (1996) simulatcd a
laterally displaced fibular fracture of 4 mm and
placed cadavcric ankles under a 500-N load in vm"·
ious static positions of ankle dorsiflexion and plantarnexion. Cutting the deltoid ligament in this
study did not result in significant lateral talar shift
or alteration in joint contact area or pressure. They
hypothesized that under static loading the talus
moves to a position of maximum congruence
within the mortisc rather than displacing laterally
with the distal fibula. Most studies agree that the
deltoid ligament and medial malleolus arc most important in resisting talar external rotation and latcral talar shift.
1
.:'
/,'
r
)
5ubtalar joint and ankle joilll inversion are ofl
difficult to separate clinically. The calcaneoribular l
ament provides stability to inversion and torsion
stresses to both the ankle and sublalar joints. Stephe
and Sarnmarco (1992) provided an inversion stre
(0
cadaveric ankles and sequentially sectioned t
anterior talofibular and calcnncofibular ligamen
They found that lip to 500/0 of the inversion observ
clinically was coming from the subwlar joint. T
structures that contribute to stability of the subia
joint afC the calcancofibular ligament, the cervic
ligament. the interosseous ligament, the lateral ta
calcaneal ligament, the ligament of ROtlvicrc, a
the extensor retinaculum (Harpel: 1991) (Fig. 9-58
Kinetics
of the
Ankle Joint
The reaction forces on the ankle joint dudng g
arc equal to or greater than the hip and knee join
respectively. The following static and dynam
analyses give an estimate of the magnitude of the
action forces acting on the ankle joint during stan
ing, walking, and running.
Ligaments of the lateral ankle and subtalar joints: 1, ant
rior talofibular ligament; 2, posterior talofibular ligame
3. calcaneofibular ligament; 4, lateral talocalcaneal liga·
ment; 5, fibulotalocalcanealligament, or ligament of Ro
viere; 6, cervical ligament; 7, ligament of the anterior ca
sule of the posterior talocalcaneal joint; '8, interosseous
ligament,
STATICS
In a static analysis of the forces acting on the all
klc joint, the magnitude of the force produced by
contraction of the gastrocnemius and the soleus
muscles through the Achilles tendon, and consequently the magnitude or the joint reaction force.
w
The Free-Body Diagram of the Foot
A, On a free-boely diagram of tile fOOl, including the
talus, the lines of clpplication for 1/1/ anel A are extended
until they intersect (Intersection point). The line of appli
can be calculated through LIse of a free-body dia-
Hon for J (dotted line) is then delermlned by connecting
gram. In the following example. the muscle force
transmitted through the Achilles tendon and the
its point of application. the tibiotalar contact point, wit
the interseCllol1 point for VV and A (Calculation Box Fig.
reaction force on the ankle joint are calculated for
1- 1). B, A tnangle of forces is constructed. Force A is 1
a subjecl slanding on liptoe on one leg. In Ihis example, the foot is considered a free-body with
three main coplanar forces acting on it: the
Ilmes body ,-,veight and force J is 2.1 times body weight
(Calculation Box FIg. 9-1-2).
ground reaction force (\V), the muscle force
lhrough the Achilles tendon (A), and the joint reaClion force on the dome of the lalus (J) (Calculalion Box 9-1).
The ground
,"
reaction force (equaling
body
weight) is applied under the forefoot and is directed upward venicallv. The Achilles force has an
unknown magnitude but a known point of application (point of insertion on the calcaneus) and a
known direction (along the Achilles tendon). The
talar dome joint reactive force has a known point
of application on the dome of the talus, but the
Illagnitude and line of direction are unknown. The
magnitude of A and J can be derived by dc:signating the forces on a free-body diagram and constructing a triangle of forces. Not surprisingly,
these forces are found to be quite large. The joint
reactive force is approximately 2.1 times body
weight. and the Achilles tendon force reaches approximately 1.2 times body weight. The great force
required for rising up on tiptoe explains why the
patient with weak gastrocnemius and soleus muscles has difficulty performing the exercise 10 times
in rapid succession. The magnitude of the ankle
joint reaction force explains why a patient with degenerative arthritis of thc ankle has pain while rising on tiptoe.
An in vilro sludv bv Wang el al. (J 996) found lhal
lhe fibula Iransmits 17% of Ihe load in the lo\\'er extremity. \Vith the ankle positioned in varus or plantarncxion, the [-Ibular load decreased. During ankle
valgus or dorsiflexion, fibular load transmission in~
creased. Cutting the distal syndesmotic ligaments
decreased fibular load transmission and incrc:ased
distal fibular migration. Cutting the interosseous
membrane had no effect on fibular load transmission. The distal syndesmotic ligamcnts arc therefore
importanl for pn;venting distal migration of the
fibula and maintaining Ilbular load.
Tibiotalar
contact point
t
ForceW
A
Calculation Box Figure
9~1-1.
~a
'\
:,\ \
~
,:
:
"
Force J ~
Force A
1.2 W
\
\
\
-
2.' W :
B
Force W
~
Force W = Ground reaction l
Force A = Muscle force thro
Achi11eslendon
Force J = Joint reaction forc
the dome of the la
Calculation Box Figure 9-1-2.
ANKLE LOAD DISTRIBUTION
The ankle has a relatively large load-bearing surface
area of II to 13 CIl1~, resulting in lower stresses
across this joint than in the knee or hip (Greenwald.
1977). The load dislribution on lhe talus is determined by ankle position and ligamentous integrity.
During weight-bearing, 77 to 90('/0 or the load is
transmitted through the tibial plafoncl to the talar
dome, with the remainder on the medial and lateral
talar facets (Calhoun et aI., 1994). As lhe loaded an-
A
B
c
kle moves into inversion, the medial taJar facet is
loaded morc. Ankle eversion increases the load on
the lateral tabH facet. The centroid of contact area
moves from posterior to anteriOl: frolll plantarOexion
to dorsiOexion, and from medial to lateral during the
motion frorn inversion to eversion. The total lalar
contact was greatest and the <.werage high pressure
was lowest in ankle dorsinexion (Calhoun et al..
1994) (Fig. 9-59).
TalaI' load distribution is also determined by ligamentous forces. Sectioning of the tibiocalcancal fascicle of the superficial deltoid ligament in a loaded
cadaver model resulted in a 430~) decrease in talar
contact area, a 30 fYo increase in peak pressures, and a
4 mm lateral shifl of the centroid (Eadl el aI., 1996),
Dynamic studies of the ankle joint are needed to appreciate the forces that act on the normal ankJe during walking and running. Stauffer et al. (1977) used
force plate, high-speed photography, radiogntphs,
and free-body calculations to determine ankle joint
compressive and shear forces. The main compressh'c
force across the normal ankle during gait is produced
-i·i.by contraction of the gastrocnemius and soleus mus"··,des. The pretibial musculature produces mild com,>;'pressive forces in early stance of <20% body weight.
'/~ compressive Force of five times body weight was
';produced in late stance by contraction of the poste,'rial' calf musculature (Fig. 9-60A). The shear force
·~:.:"reached a maximum value of 0_8 times bodv weight
"dm'ing heel-off (Fig. 9-608). Proctor and Pa~11 (1982)
also measured ankle compressive forces during gait
and found peak compressive forces of four times
body weight. In COIHrast to work by Stauffer et al.
(1977), they found substantial compressive forces
equal to body weight produced by contraction of the
anterior tibial muscle group.
The pattern of ankle joint reactive force during
gait c1iffers Wilh different walkin!! cadences (Fi!!.
,?-61). In a faster cadence, lhe paltern showed I":;'
'--J- __ ."-_ .. /
'-:C
o
E
F
Schematic demonstration of prints on pressure-sensitive
film representing high pressure contact areas on the left
talus. A, 490-N load in eversion; note lateral shift of talar
contact area. B. 490-N load in neutral version. C. 490-N
load in inversion; note medial shift of talar contact area.
0, 490-N load in 100 dorsiflexion; note anterior shift of ta
lar contact area and an increase in contact area. E. 490-N
load in 30° plantarftexion; note posterior shift of talar co
tact area. F. 980-N load in neutral; note increase in talar
contact area with increased load.
•
peak fOI-ces of three to five times body weight, on
in early slance phase and the other in late stan
phase. In the slower cadence, only onc peak force
approximately" five times bod.y weight was reach
during late slance phase (StaL\ffer et aI., 1977). Du
ing running, localized ankle forces I11ay be as hig
as 13 limes body weight (Burdett, 1982).
Effects of Shoe wear on
Foot/Ank.le Biomechanics
\¥estern society places great importance on the a
pearance of footwear, especially among wome
vVomen's footwear is designed to make the foot a
pear smaller and the leg appear longer by narrow
ing the locbox and elevating the heel. A narro
J~-"-'.""
toebox compresses the forefoot medially' and laterally, thus contributing to the development of hallux
valgus, hammer toes, and bunionettes. A study of
356 women by Frey ot al. (1993) found that 88%
of women \\'ith foot pain wore shoes that \vere on
average 1.2 em narn)\ver than their foot. Women
who \vore shoes on average 0.5 em wider than the
foot had no symptoms and less deformity. Shoes
with elevated heels increase forefoot pressure compared with standing barefoot (Snow, \Villiams, &
Holmes, 1992). A 1.9 cm heel increased forefoot
pressure by 22%, a 5 em heel increased peak pressure by 57%. and an 8.3 CIll heel increased peak
pressure b:v 76%. An elevated heel can cause pain
"n"
Ankle joint reaction force expressed in multiples of
weight in a normal ankle during the stance phase o
at two velocities. In the faster cadence, there were
peaks of three to five times body weight, one in ea
stance and one in late stance phase. In the slower c
only one peak force of approximately five times bo
weight was reached during late stance phase. Repr
with permission from Stauffer, R.N., Chao, E.Y.S., &
ster, R.L. (1977). Force and motion analysis of the n
diseased and prosthetic ankle joints. (lin Orthop, 12
Normal subjects
Preop. pIs.
,,_. Postop. pIS.
5
~ 4
'0;
;:, 3
'8
£
2
"'"
o
u.
100
A
20
40
60
under the metatarsal heads and mav also
tribute to interdigital neuroma formation.
tion of thc heel also Illa}' ovcr time res
Achilles contracture, limited ankle dorsif
and an altered gait. The amount 01" ankle joi
tion in the gait cycle decreases as heel hei
creases (Murray ct aI., 1970).
0.4
i"
:E
C> 0
'Q5 lL.
~
>.
0
D
0.2
0.2
£
'"
"
0
u.
""..,...•"..,.
0
t;:
«
0.4
0.6
0.8
'"<;;'"
'0
'iD
'"
I
Summarv
100
60
Percentage of Cycle
(stance phase)
B
A, The compressive component of the ankle joint reaction
force expressed in multiples of body weight during the
stance phase of normal walking for five normal subjects
and nine patients with joint disease before and after prosthetic ankle replacement. B, The fore-aft shear component
produced in the ankle during the stance phase of walking
for the same subjects. Reprinted with permission from Stauffer, R.N., Chao, E. Y5" & Brewster, R.L. (1977). Force and motion
analysis of the normal, diseased and prosthetic ankle joints. (fin
Orthop, 127, 789.
1 The I"oot alternates in form and functi
t\veen shock-absorbing flexible platform and
propulsive lever during different phases of t
cvcle.
2 The ankle and subtalar joint act like a m
hinge. Ankle dorsiflexion and tibial internal ro
are associated with subtalar eversion (pron
ankle plantarllcxion and tibial external rotati
associated with subtalar inversion (supinatio
3 SlIbtalar motion is screw-like and infl
the flexibility of the transverse tarsal joinl.
lar inversion locks the transverse joint and
the foot to become rigid; subtalar eversio
locks the transverse tarsal joint and allow
flexibility.
"4 The Lisfranc's joint (tarsornctt\tarsal joints) is
intrinsically stable and relatively immobile as a result of its arch-like configuration and the key-like
structure of the second tarsometatarsal joint.
15 The deltoid ligament prevents ankle ever
sion, extcTnal rotation, and lateral talar shift. It i
kc~" in maintaining the integl-it,Y of the syndes
mosis.
,~ The first metatarsophalangeal joint exhibits a
wiele range of motion, with gliding throughout Illost
or its range and jamming at full extension.
16 The fibula bears approximately one sixth of th
force exerted through the lower extremity.
<,§:,(' The medial longitudinal arch acts like both a
beam and a truss. The arch is elevated through the
windlass mechanism of the plantar fascia. The posterior tibial tendon provides c1ynarnic support to the
arch.
? Foot muscle action during standing is relatively silent, but sequential firing of both extrinsic
and intrinsic muscles is necessary to produce a normal gait patlcrn. The anterior tibial musculature
fires during early stance to slow fOOL plantarOexion
and prevent fOOl slap. The posterior cnlf musculature fires during mid- and late stance to control pro~
grcssion of the body over the foot.
~8 During barefoot standing, the heel bears 60%
or the load and lhe forefoot bears 28(Yr!. Forefoot
peak pressures occur under the second metatarsal
head.
<,9 During walking, the ccnter of pressure moves
from the posterolateral heel rapidly across the midfoot to the forefoot with peak pressures under the
second 01· third metatarsal head. At toe-ofr. the hallux bears the most pressure.
·10 The heel fat pad is specificallv designed to absorb shock dUI'ing heel strike. The plantar fascia attaches the skin of the hecl and forefoot to Ihe underlying bony and ligamcntous strUClures_
1:1 The ankle joint has muhiple axes that change
during motion. Minimal talar rotation occurs during dorsiflexion and plantarflexion.
',Tt: Ankle joint inslant centers of rotation fall
within the talus during range of motion. In 1ll0VCrncnt from plantarnexion to dorsiflexion, joint surfaces first distract, then glide and eventually jam at
the end of dorsiflexion.
13 Ankle joint stability is determined by joint congruency and ligamentous integrity. Ankle stability
increases and depends marc on articular surfacc
congnlency during weight-bearing.
14 The anterior talofibular Hnd calcancofibular
ligaments synergistically provide stability againsl
inversion during ankle motion.
JZ The distal syndesmotic ligaments preven
separation of the distal fibula and tibia and hel
lransmit force through the distal fibula on wcight
bearing.
18 Ankle joint centroid (cenler of pressure) pos
tion changes wiLh ankle flexion-cxtension and inver
sion-eversion. TalaI' surface contact is maximize
and joint pressure is minimized in dorsiOexion.
19 The rorces acting on the ankle can rise to level
exceeding five tin1es body weight during walkin
and thirteen times body weight dudng running.
20 Narrow shoes and high heels can ad\'ersel~' af
fect foot mechanics, leading to forefoot deformities
heel pain, and Achilles contracture.
REFERENCES
Adt::l:wr. R.S. (1986). Tht:' pr;ll.'tical biomechanics of running
:\111 J Sporls Jlcd. I -I. -197.
Ahu. T.K .. KilJoka. I·I.B ..
1..110, Z.I)., el at. (1997). KincmaLic
and Conl.acl c1lilracl(.'ristics of thL" first mCI:lI'.lrsoph
langt';d joint. Foot AI/kit' 111/. 18. 170.
Astion, D.J., Debnd, J.T., Otis, J.e., l.'1 <II. (1997). Motion o
Ihe hindfool after simulated arthrodesis. J 8011t' Joil1
S/lI'g. i9A. 241.
Ataarian, D.E .. t\1cCrac:kin. 1·1.1.. De\·ito. D.P., CI ill. (1985
Bionll.'chanical char:ICIl.'ristics of human ankll.' lig<lItll.'lll
Foot .·lllklc', 6. 54.
BaruNI. C.J. &. N<lpier, l.R. (1951). The axis of rol<llion at th
ankle joinl in Illall. Its influl.'llCe upon [hi.' form of Ihl.' wlu
:lIld the mobility of the fihula. J fhulf, 86, I.
Basmaji<1Jl. J.V. &: Stedo, G. (1963). The role of flHI:>l:!l.'S
<\rch supporL of Iht:: fool. J BOllI: Joint Sllf~. 45..\. 1184.
Beaudoin, A.J., Fiort::, S.~·l., Knillse. \V.R .. eL al. (1991) Effe
of isolated talocalcaneal fusion on conlact ill Ihe :Ink
and talonavicular joints. Foor AI/Uf, 12, 19.
Bbis, i\'1.~'1.. Green, \V.T., I$.: . \IHll.'rson, M. (1956). Ll.,tlgths
the growing fool. J BOllI: JOiHl S/lrg. 38:\. 988.
13ojsen-j\lolIl.'l'. F. &: LllllOI'<..'II.\:. l.. (1979). SignifiC:lIlCc of fr<
dorsifk'xion of 1!Jl.' IOt:'S in walking. ..tcUl 0,.11101' Sc:al/ll, 5
471.
6111111':U. C.. Thonl<ls, l\..A., Hakorson, T.L., l.'1 :11. (1991). Bi
l1l('c1wnical ('\'31ualioll of the anlt..'rior drawer test: Ti
contrihution of Ihe lah.'ral ankle ligamenLs. FOUl ..Il/kll'. I
389.
Bunk,lt. R.G. (1982). Forces prl.'dicll..'d :tl Ihe :tllkle durin
rulllling . .\led Sci Sporls EXn"c, Iol, 30S-310.
Bllrll:'i. W.e. II. Prakash. K., ,\debar. R.. l.'t al. (1993). Tihiol
"II'
j()inl
dynamics:
Indit.::ltions
for
lhl.' syndl.'Slllot
:'CI'<..'W-:\ l::ldaH'r study. FOOl .-I,,"-Ie, 14. 153-158.
Calhoun. J.I-I .. Eng. :\1 .. li. E. ('1 ;d. (199·0.:\ compr... hL'nsh·c
study of !1t"l..'S:iUI"I.' dislribulion in th~' :lnkk joint with inn:rsion and cn:rsion. FOOl AJlk/~ 1111. 15. 123-[33.
Cass, J.R. &. s~,tlh.':-. H. (199-0. Auld ... in stahility: In vitro
kinL'nw.tics in r...sponsl· (0 :lxi:llioad. FoOl ..\Ilk/c II/I. 15(3).
13~-I~O.
C.l\·anagh, P.R .. Morng. E., BOU!lOIl, :\.J.I\-I.. L'I a!. (1997). Thl,,'
rl'latiollship of SI;\lic foot :-tnll.:tUl"i· to dyn;ll1lic fOOl func·
lioll. J BioJ/ll.!t:I/(lIIic_.;, 30, 243-250.
Ca\"anagh. P.R .. Rodgi.:l's. rvl.\l.. &. lib()~hi. A. (1987). PressLln:.'
distribulion undt.=r s,\"mptom . frcc !"L't:! during bardont
standing. FOOl AI/kle. 7, 262.
Cawli-.'\', P.W. &. Fmncc, E.P. (1991). BiolllC'(:h:lnio; of thc latera'i ligal1\('llls of lhl..' ankle: An e\"alll:llioll of thL' effects of
<lxialload ;lIld single plane motion 011 Iig;UHI'IlI slr;lin p,lttcrns. FOOl ..\Ilkk 12. 92.
Cheng. J.CY.. LL"ung. S.S.F.. lcung, .-\.K.l... et al. (1997).
Ch'lllgt:.' of foot sizL' with wcighthe:lring. r\ sludy of 2829
childr~'n 3 to 18 YC,lr~ of age. Clill OnJwp. 3-12. 123.
Clark. T.E. (19S0J. The Pn'Hllre lJi.";lrihl/licHl UI/{!cr fh .. FOOl
Duril/g 8(/1'4001 H'alki,l,li. Uni\'l'rsity Park. f'.... nnsyh'ania:
Pcnn Sl.:lli.' Uni\'crsity.
Close. J . R. (1956), SOIllI' applications of thi.' fUIH:tion;:tl
,lnalomy of the <lnkk joint. J BOlle Jnillf Slfr~. 386, 761.
Daly, P.J., Kitaob, I·I.B .....\: Chao. E.Y.S. (1992). Pbnl:ir f:ts·
ciotomy for intnlct:tblc planlar fasciillS: Clinical !'csults
and biomt.'dl:lnil.:~\1 c\'altmtion, FOOl AI/kIt'. 13. 188.
D:t\'is, W.l-! .. Sobl:1, M .. Diarln, E.F. ..... t :tl. (1996). Gross. his .
tological. and mil.:nwascubr analolll~' ;lIld bioll1cchanicd
tesling of the spring ligamt.·111 complex. FOOl ..II/kIt' IlIl. 17.
95.
Earll. :\1., \Vaync. J.. Brodrick. C. ct 411. (1996). Contribution
of the ddlOid ligamellt to ankle joint cont,lct characteris·
tics: r\ GH!<IH'l' study, foot AIlk/e Illf. Ii, 317,
ElfllHan, H. (1960). The transverse tars.t! joint and its contro!. Clin OnllOp, 16. 41.
rick, R. (191 I). J-/wulbuch dcr '-\/Hl{OJl/il , /fI1i! .\It'cJHlllik da Ct'..
b,ke, Jena Fischer.
Frey.
Thompson, F.. Smith. J .. ~'I al. (1993). Am~'1"ican Or·
thopaedic Foot ;llld Ankle- Society wOlllen's ~hot: stirn:....
FOOf ..iI/kit:, 14. 7S-SI.
Fritz, G.R. & Pricskorn, D. (1993). Fina IIk'lat'lrsocunciform
motion: :\ radiographic and ~tatislical an~tlysis. FOOf Allklc
llIl, /6, 117.
Grecnwilld, S. (I 97i'). Unpllhlish~d d<ll;:t dtl'd in R.N. Stauffer, E.Y.S. Chao. ~\nd C. Brewster. F()l'c~' alld mOlion analysis of the lIorll\;d. diseased, and pr(lslh~'lic ;Inkle joint. C!i!1
Onhop. 127. [89.
Harper. ;\1. (1987). Deltoid ligamcllt: An t:Xpl.'l'illlClltal cvalua .
tion of funclion. FOOl Allkle, 8. 19.
Harpcr. '\'1. (1991). The lateral ligamentous support of tht.'
subtalar joint. FOOl Allklt', II, 334.
Hil-ks, J,H. (1954). Thc mechanics of the foOl 11: The plantar
aponeurosis and the arch, ) ..lllal, 88. ~5.
Hinterm<tnn. B,
Nigg. B,M. (1993). In vitro kinem.. tics of
the .. xi .. lly loaded ankle cOJnple.'\ in r('sponsl' to dorsiflcx·
ion and pl:IIH:ll'flexion. fOOl AI/kl.: 1'11. 16. 514,
Huang, C.K .. Kitaoka, 1·1.8,. An. K,N .. et :11. (1993). Biom~'·
chanical c\'aIUalion of longitudinal <In..-h sl~lhili[y. fool AIl-
c..
'*
k!c, I{ 333.
Hutton, W.C" SCUlL J.R.R" 6: Stokes, J.:\.F. (1973) The llll'chanies of thl! fool. In L. Klentrlllan (Ed.). The FoOl tIIul
Il.' Di_\orda\ (p. 41). O'\iord: BIa . : b'l'!1
li(H1S.
Sd~·I1!i"il,.' P
or
dlt' AI/kit-. l1u
Williams 6: \\·ilkin:!o.
J~lhs:-- ..\1.1-1.. KUlllllh·r. F.. &. ,\lidldsoll. J.D. (199hJ.
g~ltion into lill' bt p:lds of Ih...· sok- of th~' foo!: H~
sure s!lldi~·s. Fl)OI .. llIklt'. 13.22/.
J;dlS~ . .\1.11 .. :\-1 ichd . . l.lI'l. J.D .. Desai. P.. r..'l :11. (1~91h).
g:ltions into lhe bt p~lds of Illl' sole of I Ill' foot: :\
and histulogy, FOOl .. II/Ide. 13.133.
Kil:\Oka. I'I.B .. Lun, 70.1'.. \,\: :\n. K.N. (1997), Efl'l· ... l oi p
libia! I~'ndoll on th ... an.:h 01' the fOOL during sirnul:llcd
hearing: Bionl~'ch:tnk,d ;uwh'sis. FOil! .. \nk/,' 1111. 18,4
l\..ial'r:-.ga:IHI-:\nd~·rsl.'n. P.. \Vdh ... lund. J.O .. I!('lmig. P
{19~9J. Stahilizing dfl'Ct (If the tibiOl:ak:lll ...· al bsc
thL' ddlOid on hindfoot joillt mO\'l'lll~'lltS: :\n ~·.·qwr
~tml.'-". FOl!! .. Iukll'. /(), 30.
"-hiliI..'. K.. H:IIlSI.'Il, S.T.. c.:. \Iasqlll·kl, ".c. (1994 L C
quanlitali\"l;:' aSSn;Sllll'lll of firsl larsolll.... till,lrs.11 m
(Ill Ihe s,lgitl;d p1;Il\~' ;lnd il~ r.. . lation 10 h.dlu.'\ ,·:d
formily. [-'OOf Allkll' 1111. 15( [J, 9-13.
l.all1<lr~·all.\, L.\\·. (1971). Kin~'lIlatiL' ITh:;:lsurl'IlII,,'nb
~tlldy of hlllll:1I1 w'ilking. Bilil Prosthl'f Rl·.,. 10. I.
Lundh...-rg, ..\ .. Goldil'. I .. K:ilin. B., ~'l :11. (19~9a). Kin~·
(If IhL' :lnkkffolll ('(Illlpk-\: PI;llItarfll·,ion :Ind dl,rsi
FOol .. InUt'. 9. 194-100.
l.undbl'l"g . ...\ .. S'·l'ns:-.l.ln. O.K .. B.\"Itllld. C.. ~'I :11. (
Kinel1latics of I Ill' :lnkldfoOI cOlllplc-x,--P:lrt ~: Pr
:Illd slIpin:llio!l. rool ..\I/h!('. 9. 248.
l.undh(·r!! ...\ .. S\·.. ' llsfln. O.K .. N"'111('[1I. G.. (.[ :'d. (1%9i...'
lllati(~ 01' Ihe :lllkk,/foOI ~·()Jllpk·.\--Parl 3: lnfllll'IK
n:Il:llioll. roO! :llIkll'. V. 304.
Lundherg, A., S\"('llS~lll. O.h .. N(:llli..'th, G., l'l :11. (1989£
axis of rotation of thc :llIkk .ioint. J BOlI l ' Joilll Sll
lnmall. \'.T. (19761. Til. JO;IH_'
94-Y9,
.\l~lnn, R...\. (1982>. Biollll·t.:h:1I1ics of rUllllill~. In
plIsilllJl 011 Ihl' FO//l /llId I.l'~ ill R/IlIII;II~ Sp{/1'(S
..l,.\O
(pp.
SL Louis: C\· .\Iosh\" Co.
'\!anll. R ..-\. lI97')). Bi~md:hJni\'·s of th ... rnO!. In ..\ ..\O
(It" Orr!lolit·s: Bioll1n:lltIlJint! PrillciJJk,; tllld :l{lp/iclIt
257-266). Sl. Louis: C\', .\Io:!ob.\' Co.
.\bnn. R..-\. (1993)' Biollk'l:h:lllics of thl' foot :Illd au
SlIl";.!ay o{ llli: FpPI illld ..\/lk/t· (PP. 3-431. 51. L.ol
\Iosby Co.
:\lallkr. J.T. (19411. '\\()\"l'r1ll'IHS of thl' sllhwlar Hnd Ir;l
lars;l! joinb..··1))111 Rl'(.'. 80. 397.
Mit.:helsnn. J.D .. Clal'h'. 1-1../ .. &. Jinnah. R.I-!. (1990).
f~'cl 01' loading on libiol:dar alignlllent in (,;,Id;\\'er
fOOl .-\111:11.'. 10. 280.
\lidl~·lson. J.D. 6.: Hl.'lgl'Il\C'. S.L. Jr. (1995). Kill~'lllati(
;l.\i:tll~· loaded :Inkk. rHiIl ..\l1klt' 1111, 16,577,
:\tizl.:I. ,\I.S. (1993). Th~' role of the plallt~tr first me
l.·ul\~'ir()rll1 ligallwnt
in w('iglllhearing 011 11
Illl'tatarsai. fOOl .-\lIkle. I-l, 82.
.\I()rlon, D,J. (l935L TlI(' /lumlill Fool. liS Emlmiol/.
(1';:.'" 111111 FII//dio//al J)i...;on/as. ;'\('\\. York: COllllllhi:1
Sil.\· Pr...· ss .
.\ Illrra.,·. .\-J.P., Drought, .-\.11 .. &. Kor.\'. R.C (1964). W
paul'rlls in nnnll:ll Ill~·n. J BOllt' JOill1 51/1':-:, ,lti.-\, 33
~\Iurray, ,\I.P., Knry, R.C, 6.: Sepie, S.B. (19;0). \V'llki
terns
normal woml'n . .-II'"h I'hys ,\fed Udwhil. 51
or
Nigg, B.:'V1., Skorran, G., Frank, e.8., l't .d. (1990), Elongation
and forces of ankle ligamellts in a ph:,siological range 01
motion. Fool ..\/ll:Ie. I J, 30.
Okrud, C. &. Rosendahl. Y. (1987). Torsion-transmitting properties of the hindfoo!. Cfill On/wI', 2/4, 285.
Ohdrie-Harris, D.1 .. Reed, S.c.. &. Hedman, T.P. (1994), Dis~rllplion of the ankk syndesmosis: Biomechanical stud~' of
ligamentous reslrainb. /\rlhroscopy, 10, 558~560.
Ouzoniall, T.1. & ShL'I'l'IT, \.Lf. (1989l. In \'tIro dClL'rrnination
of midfool motion. 1"001 Allkte, !O, 140.
Pereira, D.S .. Koral, K.J., Resnick, R.B .. ct al. (1996), libiotalar contact area and pressure distribution: Th(' effect of
mortise widening and syndesmosis fixation. FOOl AI/kle
11Il, 17, 269.
Perry, J. (1983), Anatomy and biomechanics or the hindfoot.
Cfill OrtflOP, 177, 9.
Proctor, P. &: Paul. J.P. (1982). Ankle joint biomechanics. J
Bio!llcch, /5, 627.
Renstrom, P., \Venz, \1., Inc<l\'o, S., et aL (1988). Strain on
the lateral ligaments of the ankle. FOOl Ank/e, 9, 59.
Sammarco, G.J. (1980). Biomechanics of the fool. In \'.1-1.
Frankel &: \-1. Nordin (Eds.), Basic Biolllcchallics of Ihe
AllisclIloskdCf(/! Systelll (2nd ed., pp. 193-2(9). Philadelphia: Lea &. Febiger.
Sammarco, G.J. (1995). Peroneus longus tendon tears: Acute
and chronic. Foot Ankh, lilt. /6(51,245-253.
Sa tTl lTIan:o , G.J .. Burnstein, A.H., &. Frankel. V.I-I. (1973). Biomechanics of the ankle: A kinematic study. O/"lhop Cli/I
Norlh Alii, -t, 75.
Sarrafian, S.K. (1987), Functional characteristics of the foot
and plantar aponeurosis under tibiotalar loading. FOOl .'111l:lc, 8, 4.
Sal'l'afian, S.K. (1993a}. FUllctional anatomy of the foot and
ankle. In An(ltomy or the Foo! ami Anl:lc (pp. 474-602l.
Philadelphia: Lippincott.
Sarnlfian, S.K. (1993b). Retaining systems and compartments. In A/UlIOIllY
the Fool (lilt! Auk/e (pp. 137-149).
Philadelphia: Lippincott.
Scranton, P.E., ;V\c,\'laster, J.I-I., &. Kelly, E. (1976). Dynamic
fibular function: A new concept. Clill OrfflOp, 118,76-81.
or
SherefT, ,\l.J., Ikjahi, F.J .. &. Kummer, F.J. (1986). Kinern
or the first met~llarsophabngeal joint. .I BOlle Joint S
68..1, 392.
Silver, R.I.., de la G~lrz~l, .I., l\: Rang, M. (1985). The my
llluscle balance: i\ stud~ of 1"t.'lati\'e strengths ~lnd ex
sions about the foot and ankle. J BOlle Joiu! Surg,
432.
Snow, R.E., \Villiams, K.R., & Holmes, G.B. Jr (1992). Th
reets of wearing high heeled shoes on pedal pressure
women. Foo! Allkle, /3. 85-92.
So~\nll's, R.\\'. (1985), Foot pressures during gait. J 13io
Eng, 7, 120-126.
Stauffer, R.N., Chao, E.'{S., &: Brewster, R.I.. (1977). F
and !'ll(Jti(lll ~lnal.\'sis of till: n()1'mal, disl'<lsl:d and
thetic ankle joints. Clin Orthop, /27, 189.
Stephens, :\1.\1. & Sammarco, G.J. (1992). The stabili
role of the lateral ligalllt.'nt complex around the ankle
subtalar joints. FOOl AI/klc, 13, 130.
Stiehl, J.B., Skrade, D.A., Needleman. R.I.., et al. ([ 993)
fect of a.'\i~ll load and ankle position on ankle stabili
Ort/lOp 7/"al/l1l(/, 7, 72-77.
Stormont. D.\I., :\10rrey, B.F.. An, K.N., et al. (1985). St
ity of the loaded ankle. Relation between articula
straint and prinwr.\' and sec()[ldar~' restrainis . ..1111 J S
Met!, 13, 295.
Thordarson, D.B.. Schmotzer, H . Chon, .I .. et al. (1995)
namic support of the human longitudinal arch. Clill
lhop. 316, 165.
Wang, O.W., Whiltle, .\1.. Cunningham . .I .. et al. (1996). F
~Hld its ligaments in load transmission and ankle joint
bilit.\". Clill Ort/lOp. 330. 261.
Wanin:nhaus, A. & Pn:Llcrklieber, \1. (1989), First
sometatarsal joint: Anatomical and biomeehanical s
FOOl Allkfe, 9, 153.
\Vaters, R.I.., Hislop, I-U., Perry, .I., et al. (1978). Energc
Application of the study and m,Hl~lg('mcnt of locom
disabilities. Orlhop Clill North Am, 9, 351.
\Vright. D.G" Desai, S.\1.. 6: Henderson, \V.H. (19M), A
of the sublalar and ankle joilll comrle.'\ during the st
phase of walking . .1 130111.' Joil/I Sur,!!" 46.'1, 361.
Biomechanics of the
lumbar Spine
Margareta Nordin, Shira Schecter Weiner adapted fro
Margareta Lin
Introduction
The Motion Segment: The Functional Unit of the Spine
The Anterior Portion of the Motion Segment
The Posterior Portion of the Motion Segment
The Ligaments of the Spine
Kinematics
Segmental Motion of the Spine
Range of lvIotion
Surface Joint Motion
Functional Motion of the Spine
The lvIuscles
Flexion and Extension
Lateral Flexion and Rotation
Pelvic Motion
Kinetics
Statics and Dynamics
Statics
Loading of the Spine Dunng Standing
Comparative loads on the Lumbar Spine During Standing, Sittin
and Reclining
Static Loads on the Lumbar Spine During Lifting
Dynamics
Walking
Exercises
lvIechanical Stability of the Lumbar Spine
Intra·Abdominal Pressure
Trunk Muscle Co-Contraction
External Stabilization
Summary
References
Introduction
The human spine is a complex structure whose
principal functions are to protect the spinal cord
and transfer loads from the head and tnmk to the
pelvis. Each of the 24 vertebrae articulates with the
adjacent ones to permit motion in three planes. The
spine gains stability from the intervertebral discs
and from the surrounding ligaments and muscles;
the discs and ligaments provide intrinsic stability
and the muscles provide extrinsie support,
This chapter describes the basic characteristics
of the various structures of the spine and the interaction of these structures during norrnal spine function. Kinematics and kinetics of the spine afC also
examined. The discussion of kinematics covers both
the thoracic and lumbar spine. but that of kinetics
involves only' the lumbar spine because it is subjected to significantly' greater loads than is the rest
of the spine and has received more attention clini·
cally' and experimentally. The information in the
chapter has been selected to provide an understanding of some fundamental aspects of lumbar spine
biomechanics that can be put to practical use.
11
12
13
14
2
3
10
4
5
6
8
7
Posterior portion ~,_,-,~ Anterior portion
Schematic representation of a motion segment in the lum
bar spine (sagittal view), Anterior portion: 1, posterior lo
gitudinal ligament; 2, anterior longitudinal ligament; 3,
vertebral body; 4, cartilaginous end plate; 5, intervertebr
disc; 6, intervertebral foramen with nerve root. Posterior
portion: 7, ligamentum flavum; 8, spinous process; 9, inte
vertebral joint formed by the superior and inferior facets
(the capsular ligament is not shown); 10, supraspinous lig
ment; 11, interspinous ligament; 12. transverse process (t
intertransverse ligament is not shown); 13, arch; 14, verte
bral canal (the spinal cord is not depicted).
Molion
segment
The Motion Segment: The
Functional Unit of the Spine
I
~----I
I
1
I!l
Anteroposterior (A) and lateral (B) roentgenograms of the
lumbar spine. One motion segment, the functional unit of
the spine, is indicated.
The functional unit of the spine, the motion segmen
consists of two vertebrae and their intervening so
tissues (Fig, 10- I). The anterior pOl,tion of the se
ment is C0111posed of two superimposed intervert
bral bodies, the intervertebral disc. and the longit
dinal ligaments (Fig. 10-2). The c9ITcspondi
vertebral arches, the intervertebral joints formed
the facets, thc transverse and spinolls processes, an
various ligaments make up the postcrior portion. T
25
Anterior
Vertebral body
Iliopsoas
muscle
Arch
Spinal canal
Transverse
with cord
process
Interspinous
Facet
ligament
Spinous
process
Erector
spinae muscle
Posterior
The nucleus pulposus Iic's directl~' in the cente
all discs e\cept tllOSL' in tilL' IUlllhar segments, wh
it has a slightl~, posterior position. This inner m
is surrounded b,\' ;:\ tough outer covering, the ann
lus fihrosus, composed 01" fibr()cani];:I!~:e. The cri
cross arrangcrnent 01" the coarse collagen fiber b
dles within till' fibrocartilage all<)ws the annu
fibrosus to withstand high lk'!l{ling and torsicJ
loads (se(' Fig, II-II). Discs with annular tears d
pia." increased [,(ltational 1T'IOlllents during k,ad
compared with IHlndegenerated discs (Haughton
aI., 2000). The end phlte, composed of hyaline ca
lage, separ(a~s the disc from the vertebr<:11 bod~' (F
10-2). The disc composition is simihlr to that
ticular cartilage, described in detail in Chapter 3
During dail.v activities, the disc is loaded in a co
pIc.\: manner and is usually subjected to a combi
tion ()I" comp]'ession, bending, and lo]·sioll. Fle.\:i
extension, and lateral lle.\:ion or the spine produ
rnainl~' tensile ,Ind compressi\'(' stresses ill the d
\\'hereas rotation produces mainly shear stress.
\\-'hen <:1 motion segment is tr<:lllsected vLTtica
the llucleus pulposus 01" the disc protrudcs, indic
ing that it is under pressure. !\iIcasurement (If the
tracliscal prcssure in normal and slightl." degenera
cadavcr lurnbar nuclei pulposi has shown an intr
or
Transverse section of a motion segment at the L4 level
viewed by computed tomography. The vertebral body,
arch, spinal canal with spinal cord, and transverse
processes are clearly seen. The view is taken at a level that
depicts only the tip of the spinous process, with the interspinous ligament visible between the spinous process and
the facets of the intervertebral joints. Directly anterior to
I
..
the transverse processes and adjacent to the vertebral
body are the iliopsoas muscles. Posterior to the vertebral
body, the erector spinae muscles can be seen.
_----------
arches and vertebral bodies form the vertebral canal,
which protects the spinal cord (Fig. 10-3).
THE ANTERIOR PORTION
OF THE MOTION SEGMENT
The vertebral bodies are designed to bear mainly
compressive loads and the.\,' arc progressively larger
caudally as the superimposed weight of the upper
body increases. The vertebral bodies in the lurnbar
region are thicker and wider than those in the thoracic and cervical regions; their greater size allows
thenl to sustain the larger toads to which the lumbar spine is subjected.
The intervertebral disc, \vhich bears and distributes loads and restrains excessive motion, is of great
mechanical and functional importance. It is well
suited for its dual role because of its location between the vertebrae and because of the unique composition of its inner and outer structures. The inner
portion of the disc, the nucleus pulposus, is a gelatinous mass. Rich in hydrophilic (\vater-binding)
glycosaminoglycans in the young adult, it diminishes in glycosaminoglycan content with age and
becomes progressively less hydrated (Urban &
McMullin, 1985).
Distribution of stress in a cross·section of a lumbar disc
der compressive loading. The compressive stress is highe
in the nucleus pulposus, 1.5 times the externally applied
load (F) per unit area. By contrast, the compressive stres
on the annulus fibrosus is only approximately 0.5 times
externally applied load. This part of the disc bears pred
inantly tensile stress, which is four to five times greater
than the externally applied load per unit area. Adapted
,vith permission from Nachemson, A. (!975j, Tmvards a berr
understanding of back pain: A reviev'/ of the mechanicS of th
lumbar disc RheuI"T1illol Rehabi!, j'l. 129
sic pressure in the unloaded disc of approximately
\0 N per square centimeter (Nachemson, 1960). This
intrinsic pressure, or pre-stress, in the disc results
from forces exerted by the longitudinal ligaments
and the ligamentum flavum. During loading of the
spine. the nucleus pulposus acts hydrostatically
(Nachemson. t 960), allowing a unifonn distribution
of pressure throughout the disc; hence, the entire
disc serves a hydrostatic function in the motion segment, acting as a clishion between the vertebral bodies to SLare energy and distribute loads.
In a disc loaded in compression. the pressure is
approximately 1.5 times the externallyay)plicd load
per unit area. Because the nuclear material is only
slightly compressible, a compressive load makes the
disc bulge laterally; circumferential tensile stress is
slIst"ined by the annular fibers. In the lumbar spine
the tensile stress in the posteriol- part of the annulus
fibrosus has been estimated to be four LO five times
the applied axial compressive load (Galante, 1967;
Nachemson, 1960, 1963) (Fig. 10-4). The tensile
stl·ess in the annulus fibroslis in the thoracic spine
less than that in the lurnbar spine because of diffe
ences in disc geometry. The higher rario or disc d
ameter to height in the thoracic discs reduces the ci
cumferential stress in these discs (Kulak et aI., 1975
Degeneration of a disc reduces its protcoglyca
content and thus its hydrophilic capacity (Fig. 10A-C). As the disc becomes less hydrated, its elasti
ity and its ability to store energy and distribu
loads gradually decrease; these changes make th
disc(s) more vulnerable to stl-csses.
THE POSTERIOR PORTION
OF THE MOTION SEGMENT
The posterior portion of the Illotion segment guid
its movement. The t~'pe of motion possible at an
IC\'e1 of the spine is determined b~" the oriental ion
the facets of the inlervenebral joints to the tran
verse and frontal planes. This orientation chang
throughollt the spine.
Human intervertebral disc composed of an inner gelatinous
mass, the nucleus pulposus (NP), and a tough outer covering,
the annulus fibrosus (AF). A, Normal young disc. The gelatino
nucleus pulposus is 80 to 88% water content (reprinted with pe
mission from Gower. WE. &Pedrini, V. 1969. J Bone Joint Surg, 51
1154). Age-related variations in protein-polysaccharides from
human nucleus pulposus. annulus fibrosus, and costal canilag
are easy to distingUish from the firmer annulus fjbrosus. B. No
mal mid-age disc. The nucleus pulposus has lost water conten
a normal degenerative process. The fibers on the posterior pa
of the annulus have sustained excessive stress. C, Severely degenerated disc. The nucleus pulposus has become dehydrated
and has lost its gel-like character. The boundary between the
nucleus and the annulus is difficult to distinguish because the
degree of hydration is now about the same in both structures
~!!II!JIIII!~~~~~~~~~~-----'---~~;~"_'.<"~
-~_.,__
;- ... _~-<C-'t'"
Except for the facets of the two uppermost cervical vertebrae (CI and C2), which arc parallel to the
transverse plane, the facets of the cervical inten.1crtcbral joints are oriented at a 45° angle to the transverse plane and arc parallel to the frontal plane (Fig.
10-6A). This alignment of the joints of C3 to C7 allows flexion, extension, lateral flexion, and rotation.
The facets of the thoracic joints arc oriented at a 60
angle to the transverse plane and at a 20° angle to
the frontal plane (Fig. 10-68); this orientation allows
lateral flexion, rotation, and some flexion and extension. In the lurnbar region. the facets arc oriented at
right angles to the transverse plane and at a 45°
angle to the frontal plane (Fig. 10·6C) (White &
Panjabi. 1978). This alignment allows Oexion, extension, and lateral Oexion, but almost no rotation. The
lumbosacral joints differ from the other lumbar intervertebral joints in that the oblique orientation of
the facets allows appreciable rotation (Lumsden &
Morris, 1968). The above-cited values for facet orientation are only approximations, as considerable
variation is found within and among individuals.
The facets guide movement of the Illotion segment
and have a load-bearing function. Load-sharing between the facets and the disc varies with the position
of the spine. The loads on the facets arc greatest (approximately 30% of the lotal load) when the spine is
hyperextended (King et aI., 1975). Because the facets
arc not the primary SUPPOI1 structure in extension, if
total compromise of these joints occurs, an alternate
path of loading is established. This path involves the
transfer of axial loads to the annulus and anterior
longitudinal ligament as a way of supporting the
spine (Haher et aI., !994). High loading of the facels
is also present during forward bending, coupled with
rotation (EI-Bohy & King, 1986). The vertebral
arches and intervertebral joints play an important
role in resisting shear forces. This function is demonstrated by the fact that patients with deranged arches
or defective joints (e.g., from spondylolysis and listhesis) are at increased risk for fOl'\vard displacement
of tbe vertebral body (Adams & Hutton, 1983; Mi11er
et aI., 1983) (Case Study 10·1). The transverse and
spinolls processes seniC as sites of attachment for the
spinal muscles, whose activity initiates spine motion
and provides extrinsic stability.
0
THE LIGAMENTS OF THE SPINE
,,-
The ligamentous structures surrounding the spine
contribute to its intrinsic stability (Fig. 10-2). All
spine ligaments except the ligamentum Ibvulll have
;;;:
~_Q-
'
..
;
-
a high collagen content, which limits tht.:ir ext
bility during spine motion. The ligalllL'llllllll fla
which connects two adjaCL'1l1 \'(,I"lcbral arches l
(lldinall~·, is an exception. having a lal'gc percen
of dastin. The clasticit.Y of (his ligamclll <.lIlows
contract during c.'\h..:nsion of the spine and 10
gate during nc.'\ion. Even when the spine is in a
tral position, the ligamentum flavul11 is uncleI'
Slant tension as a result of its elastic prope
Because it is located at a distance from the cent
motion in the disc, il pre-stresses the disc; th
along with the longillldinal ligaments, it create
irllradiscal pressure and thus helps provide intlsupport to the spine (Nachemson &. EV<.H1s,
Rolander, 1966). Research suggests that with de
erative changes such as spond~"lolisthesis, tra
spurs, and disc degeneration, which rna}' lead t
stability, altered mechanical stress \vill increas
load the ligamentum fluvul11 and callSt: hypertr
(FlIkllyama el aI., 1995).
The amount of strain on the various ligam
differs with the type of motion of the spine, D
llc.'\ion, the interspinous ligaments arc subject
the greatest strain, followed by' the capsular
ments and the ligamcnturn flavum. During e
sion, the anterior longitudinal ligarnent bear
greatest strain. During latcral fle:don, the contr
eral lransverse ligament sustains the hig
strains. followed by the ligament nuvulll and
capsular ligaments. The capsular ligaments o
facet joints bear the most strain during rot
(Panjabi et aI., 1982).
Kinematics
AClive mol ion of the spine as in any joint is
duced b.'/ the coordinated interaction of n
and muscles. Agonistic n1uscles (prime mo
initiate and carry out motion and antagon
muscles control and modif." the motion, whil
contraction or both groups stabilizes the s
The range of motion differs at various levels o
spine and depends on the orientation of the f
of the intervertebraljoinls (Fig. 10-6). Motio
tween two vertebrae is small and docs not o
independently; all spine movements involve
combined action of several motion segments.
skeletal structures that influence motion o
lrunk are the rib cage, which limits thoracic
lion. and the pelvis, which augments trunk m
ments by tilting.
Cervical
(C3-C7)
A
Thoracic
B
I
,,
Lumbar
I
c
Orientation of
the facets to
the transverse plane
Orientation of the facets of the intervertebral joints (approximate values). Reprinted with permission from White, A.A.
& Panjabi, M.N. (1978). Clinical Biomechanics of the Spine.
Philadelphia: J,8. Lippincorr. A. In the lower cervical spine, the
facets are oriented at a
4S~
angle to the transverse plane
Orientation of
the facets to
the fronlal plane
and are parallel to the frontal plane. B, The facets of the
thoracic spine are oriented at a 60° angle to the transvers
plane and at a 20° angle to the frontal plane. C, The facet
of the lumbar spine are oriented at a 90° angle to the tran
verse plane and at a 45° angle to the frontal plane.
~
,'Sponaylolisthesis: Anterior Slippage of One Vertebra in Relation
to the Vertebra Below It
'A
30-year·old male gymnast complains of severe back pain
that lies betyveen the inferior and superior facets). This bilat·
periods of strenuous training and the symptoms decrease
eral defect leads to an (lnterior displacement of the vertebra
~ith rest or restriction of activity, After a careful examination
lS onto S1. As the l5 vertebra begins 10 slip forvllard, the cen
b'ya specialist and
ter of gravity of the body is displaced anteriody. To compen-
MRI films. a diagnosis was made of
spondylolisthesis at the level L5-S1 (Case Study Fig, 10-1-1),
a disease continuum. the abnormal forces placed on the inter
Case Study Figure 10-1-1.
vertebral disc leads to herniation into the neural foramina,
producing moderate stenosis. of both l5-S I nerve rootS.
Case Study Figure 10-1-2.
SEGMENTAL MOTION OF THE SPINE
The vertebrae have six degrees of freedom: rotation
about and translation along a transverse, a sagittal.
and a longitudinal axis. The motion produced during nexion. extension, lateral nexion, and axial rotation of the spine is a complex combined motion resulting fTom simultaneous rotation and translation.
~:"
pan of the trunk is displaced backward. Because thi) is
upper
(Case Study Fig, 10-1-2),
-' Physiological loads during repeated flexion-extension mo-
~,.
sate, the lumbar spine above the lesion hyperextends and the
with concurrent bilateral pars interarticularis defects of L5
tiOn, of the lumbar spine caused a fatigue fracture of the pars
,i;
•t"
interarticularis (aspect of the postenor arch of the vertebra
'''''..' , with radiation to both legs. The pain is associated with
Range of Motion
Various investigations using autopsy material or
radiographic measurements in vivo have shown di~
vergent values for the range of motion of individual motion segments, but there is agreement on
the relative amount of motion at different leve
the spine. Representative values from \Vhite
Panjabi (1978) are presented in Figure 10-7 1
lc)\v a comparison of motion at various levels o
thoracic and lumbar spine. (Representative v
for motion in the cendcal spine are included
comparison.)
lnvestigations of the thoracic and lumbar s
show that the range of l'lexion and extension i
proximately 4° in each of the upper thonlcic mo
segments, approximately 6° in the mid thoraci
gion, and approximately 12° in the two lower
racic segments. This range progressively incre
in the lumbar Illotion segments. reaching a m
mum of 20' al Ihe lumbosacral level.
o
I
10'
'
20
I
0'
I
10
'
20
I
0'
I
10'
'
OC- Cl
Cl - 2
2 -3
3-4
4 -5
5 -6
20"
I
'-'47'
6-7
C7 - Tl
Tl - 2
2-3
3-4
4-5
5-6
6-7
7-8
8-9
9 - 10
10 - 11
11 - 12
T12 - Ll
L1 - 2
2-3
3-4
4-5
L5 - 51
Flexionextension
Lateral
flexion
Rotation
A composite of representative values for type and range of motion at different levels of
the spine. Reprinted with permission from White, AA & Panjabi, M.N (978). Clinical Biomechanics of the Spine. Philadelphia: J.B. Lippincorr.
Lateral flexion shows the greatest range in each
of the lower thoracic segments, reaching 8 to go. In
the upper thoracic segments, the range is uniformly
6°. Six degrees of lateral flexion is also found in each
of all lumbar segments except the lumbosacral segment, \vhich demonstrates only 3° of 111otion.
Rotation is greatest in the upper segments of the
thoracic spine, \vhere the range is 9°, The range of
rotation progressively decreases caudally, reaching
2° in the lower segments of the lumbar spine. It then
increases to 5° in the lumbosacral segment.
Surface Joint Motion
i\''Iotion between the surfaces of two adjacent vertebrae during flexion-extension or lateral flexion may
be analyzed b:v means of the instant center method
of Reuleaux. The procedure is essentially the sanle
as that described for the cc,·vical spine in Chapter
11 (see Figs. 11-18 and 11-19). The instantaneous
center of flexion-extension and lateral flexion in
motion segment of the lumbar spine lies within t
disc under normal conditions (Fig. 10-8A) (Cosse
et aI., 1971; Rolandel~ 1966). With abnormal conc
tions such as pronounced disc degeneration, t
instantaneous center pathway will be altered (F
10-8B) (Gertzbein et aI., 1985; Reichmann et aI., 197
FUNCTIONAL MOTION OF THE SPINE
Because of its complexity,') the motion of a single nl
tion segment is difficult to measure clinically. A
proximate values for the normal functional range
motion of the spine can be given. Variations amo
individuals are large and show a Gaussian distrib
lion in the three planes. The range of motion
strongly age~dependent, decreasing ·b~v approx
mately 30% from youth to old age, although \vith a
ing, loss of range of motion is noted in flexion a
lateral bending while axial rotation motion is mai
I
Normal
G-J
iJ
A
~
rained with cvidcnc~ or increased coupled m
(iV!cGill et aI., 1999). Differences have also
noted between the sexes: men have greater m
in flexion and extension whereas women arc
mobile in later,,1 Oexion (Bie.-ing-Sorensen,
Moll & Wright, 1971). Loss of range of motion
lumbar and/or thoracic spine is compensat
mainly by motion in the cervical spine and hi
Moderate disc
degeneration
J
~
(. '
B
~
"-'",7
L5
THE MUSCLES
Instant center pathway for a normal cadaver spine (A) and a
cadaver spine with moderate disc degeneration (B). Instant
centers were determined for 3~ intervals of motion from
maximum extension to maximum flexion. In the normal
spine, all instant <enters fell within a small area in the disc.
In the degenerated spine. the centers were displaced. and
hence the surface motion was abnormal. Reprinted wirh per·
mission from Gerr£bein, S.D., er al. (1985). Centrode patterns and
segmental instability in degenerative disc (Jisease. Spine, 10,257.
The spinal muscles can be divided into flexo
extensors. The main flexors arc the abdomina
cles (the reClus abdominis muscles, the intern
external oblique muscles. and the transver
dominallllllsclc) and the psoas muscles. In g
muscles anterior to the vertebral column act
ors. The main extensors arc the ercCLOr spina
cles. the multifidus muscles. and the intcnra
sarii muscles auached to the posterior eleme
general, the muscles posterior to the ve
columns act as extensors (Fig. 10-9). The ex
Anterior
Rectus abdominis muscle
Transversalis fascia
Transversalis
Internal oblique
Vertebral body
External oblique
Lumbar fascia:
Anterior
Middle
Posterior
Quadratus
lumborum
muscle
Erector
spinae muscle
Posterior
An MRI transverse cross-section of the body at the l4 level of a normal adult human spine. The major trunk ffius<les are
(R, right; l, left). By courtesy (rom Ali SfJeikzahed, PhD., Hospilal for Joint Diseases, Mr. Sinai, NYU Health, New York, NY, USA.
muscles bridge between each vertebrae and motion
segment as well as over several vertebrae and motion segments. \Vhen extensor muscles contract
symmetrically, extension is produced. "'!hen right
and left side flexors and extensor muscles contract
asymmetrically, lateral bending or t\visting of the
spine is produced (Andersson & Lavender, 1997).
Flexion and Extension
During unloaded flexion-extension range of motion,
the first 50 to 60° of spine flexion occurs in the lumbar spine, rnainly in the lower motion segments
(Carlsoo, 1961; Farfan, 1975). Tilting the pelvis forward allows for further flexion. During lifting and
lowering a (oad, this rh)rthm occurs simultaneously,
although a greater separation of these movements is
noted during lifting than during lowering (Nelson et
aL, 1995) The thoracic spine contributes little to Forward flexion of the entire spinal column because of
the oblique orientation of the facets (Figs. 10~6 and
10-7), the nearly vertical orientation of the spinolls
processes, and the limitation of motion imposed by
the rib cage.
Flexion is initiated by the abdominal muscles and
the vertebral portion of the psoas muscle (Anders~
son & lavendel~ 1997; Basmajian & Deluca, 1985).
The weight of the upper body produces further flexion, which is controlled by the gradually increasing activity of the erector spinae muscles as the
forward-bending moment acting on the spine increases. The posterior hip muscles are active in con~
trolling the forward tilting of the pelvis as the spine
is flexed (Carls!j", 1961). It has long been accepted
that in full flexion, the erector spinae muscles become inactive once they are fully stretched. In this
position, the fonvard bending moment was counU:Tacted passively by these muscles and by the
posterior ligaments, \vhich are initially slack but become taut at this point because the spine has
fully elongated (Farfan, 1975). This silencing of the
erector spinae muscles is known as the flexionrelaxation phenomenon (Allen, 1948; Andersson &
Lavender, 1997; Floyd & Silver, 1955; Morris et a!.,
962). However, Andersson et al. (1996), using wire
CICCII'()('"'' inserted in the trunk extensor muscles
~l1icl("cI by ultrasound or MRI, showed that in the
flexed position the superficial erector spinae
muscles relax, while the quadratus lumborum and
lateral lumbar erector spinae muscles become
'''''V<lIeu (Fig. 10-10). In forced flexion, the superfiextensor muscles become re~activated. From
l
Electromyography of the quadratus lumborum (QL) and
erector spinae superficial (£5-5) and deep (£S-d) muscles
Wire electrodes were inserted in QL and ES-d; surface e
trodes were used for ES-s. Five positions (a-e) of trunk f
ion are depicted. In full nonforced trunk flexion (e), the
ES-s activity is silent; however, the ES-d and QL are very
tive to counterbalance the trunk flexion movement. (Co
tesy of Eva Andersson, MD Ph.D, Karofinska Institute. Stockholm, S~veden.)
full flexion to upright positioning of the trunk,
pelvis tilts backward and the spine then exten
The sequence of muscular activity is reversed. T
gluteus maximus comes into action early toget
with the hamstrings and initiates extension by p
terior rotation of the pelvis. The paraspinal musc
then become activated and increase their activ
until the movement is completed (Andersson
Lavender~ 1997).
Some studies have shown that the concentric
ertion performed by the muscles involved in rais
the trunk is greater than the eccentric exertion p
formed by.' the muscles involved in lowering
trunk (de Looze et a!., 1993; Friedebolcl, 19
Joseph, 1960). However, this finding has been c
tradicted in several studies (Reid & Costigan, 19
Marras & Mirka, 1992). Creswell ancl Thortens
(1994) support the finding that less electrom
graphic (EMG) activity is noted during eccentric
tivity, as in lowering, despite high levels of fo
generated. The compressive load of the sp
caused by the muscle exertion produced by low
ing the trunk with a load or resistance can appro
the spinal tolerance limits, putting the back
greater risk for injury (Davis et aI., 1998).
When the trunk is hyperextended from the
right position, the extensor rnuscles are active d
ing the initial phase. This initial burst of activity
creases during funher extension, and the abdominal
muscles bccorne active to control and modify the
motion. In extreme or forced extension. extcnsor activity is again required (Floyd & Silver. 1955).
joints function mainlv as shock nbsorbers a
important in protecting the intervertebral
(Wilder et al.. 1980).
When loaded in vitro. the Sljoint exhibits
dimensional movement with joint opening r
ranging from 0.5 to 1.2 and sacrum an
0
Lateral Flexion and Rotation
During lateral Oexion of the trunk, motion may predominate in either the thoracic or the lumbar spine.
In the thoracic spine, the facet orientation allows for
Imeral Oexion. but the rib cage rcstricts it (to vmying
degrees in different people); in the lumbar spine. the
wedge-shaped spaces between the intervertcbral
joint surfaces show variations during this motion
(Reichmann. 1971). The spinotransversal and transversospinal systems of the erector spinae muscles
and the abdominal muscles arc active during lateral
I1cxion; ipsilmeral COnlractions of these muscles initiate the motion and contralateral contractions modify it (Fig. 10-11) (Andersson.& Lavender. 1997).
Significant axial rotation occurs at the thoracic
and lurnbosacral levels but is limited at other le\'els
of the lumbar spine, being restricted by the vertical
orientation of the facets (Fig. 10-6C). In the thoracic
region, rotation is consistently associated with lat·
end flexion. During lhis coupled motion. which is
most marked in the upper thoracic region, the vertebral bodies generally rotate toward the concavity
of the lateral curvc of the spine (White. 1969). Coupling of rotation and lateral nexion also takes place
in the lumbar spine. with the vertebral bodies rotating toward the convexity of the clIlve (Miles &
Sullivan. 1961). During axial rotation. the back and
abdominal muscles are active on both sides of
the spine, as both ipsilateral and contralateral muscles cooperate to produce this movement. High
coactivation has been measured for axial rotation
(Lavender et al.. 1992; Pope et aI., 1986).
Pelvic Motion
Functional trunk movements not only involve a
combined motion of different parts of the spine but
also require the cooperation of the pelvis because
pelvic Inotion is essential for increasing the range of
functional motion of the trunk. The relationship betwecn pelvic movements and spinal motion is generally analyzed in terms of motion of the lumbosacral joints. the hip joints. Ol' both (Fig. 10-12).
Load transfer From the spine to the pelvis occurs
through the sacroiliac (SI) joint. Biomechanical
analysis of the sacroiliac joints suggests that these
Left
0
L3
0
L5
80
60
40
20
0
20
VFRA (pV)
\( 1\
40
60
80
\\
FIG. 10-11
------
Example of eleetromyographic activity of the erect
spinae muscles collected with surface electrodes du
side-bending of the trunk. The figure illustrates tru
bending to the right and muscle activity at the ll,
lS level of the lumbar spine. Substantial contralate
cle activity (left) of the erector spinae muscles is re
when bending to the right to maintain equilibrium
duced wirh permission from Andersson. G.BL Orrengr
Nachemson. A. (1977). fIHradiscaf pressure. intra-abdom
pressure and myoelectric back muscle dcrivily related to
and foading. Clin OrtllOp, 129. 156,
lIsed. This involves measuring the m.voelectric aCli
it~,. of the tnlllk muscles and correlating this activit
with calculc:tlcd values for muscle contractio
forces. The values obtained correlate well with thos
,I,
{
',.
The pelvic ring with its linkage to the spine and the lower
extremities. The antero-posterior view of these structures
on film gives a hint of the irregular shape of the sacroiliac
joint 5urlaces. but an oblique projection is required for an
accurate view of the joints.
, ,,
,
posterior rotation ranging from 0.3 to 0.6°; translation ranged from 0.5 to 0.9 mm (\Vang & Dumas,
1998). In vivo analysis of the 51 joint utilizing
roentgen slereophotogrammetr~!shows joint rotation mean at 2.5 0 and translation mean at 0.7 mm
with no differences between symptomatic and
asymptomatic joints (S,uresson et aI" 1989),
Muscle forces acting on the 51 joint have a stabiliZing effect. helping to attenuate the high stress
pelvic loading (Dalstra & Huiskes. 1995).
or
obtained through inLradiscal pressure measuremen
and can thereFore be lIsed to predict the loads on th
spine (Andersson & Lavendcr, 1997; bnengren
aI., 1981; Schuhz et aI., 1982).
Another method is the lise of a mathcnHuic
model for force estimation that allows the loads o
the lumbar spine and (he contraction forces in th
trunk muscles to be calculated for various phy'sic
activities. The models are useful as predictors o
load, for load sharing analysis under different con
ditions, to simulate loads, and in spine prosthet
and instrumentation design. The precision of th
model depcnds on the assumption L1sed for the ca
culations. Two categories of models currently use
arc the EiV1G-c1l'iven model based on c!cctromyo
graphic trunk mllscle rccordings and the morc tr
ditional biomcchanical model based on trunk mo
mcnts and [OI'ces (Chaffin & Anderson, 199
Lavender et al.. 1992; Marras & Granata, 199
Sheikhzadeh, [1997]. The eflixi
pllre alld COl
billed loading OIl the recruitment pattern
ten s
lecled Intlll< IIlllscles, Unpublished doctoral thesi
New York University. New York).
or
or
STATICS AND DYNAMICS
In the following section, static loads on the lumb
spine are examined for common postures such a
standing and silting and also for Iif1ing. a commo
activity involving external loads. [n the final sectio
the dynamic loads on the lumbar spine during walk
ing and common strengthening exercises for th
back and abdominal muscles arc discllssed.
Statics
Kinetics
Loads on the spine are produced primarily by body
weight. muscle activity, pre-stress exerted by the ligaments, and externally applied loads. Simplified calculations of the loads at various levels of the spine
Can be made with the usc of the [Tee-body technique
for coplanar forces. Direct information regarding
loads on the spinc.,(ll the level of individual intervertebral discs can bq/~btained by measuring the presSure within the d(scs both in vitro and in vivo. Because this mel~od
is too complex for general
;
application. a semidircclmcasuring method is often
'.".....
.'
.•.. '
The spine can be consich.:red as a modified elast
rod because of the nexibility of thc spinal colum
the shock~absorbing behavior of the discs and ve
tebrae. the stabilizing function of the longitudin
ligaments, and the elasticity of the ligamen
flavum. The two curvatures of the spine in the sagi
tal plane-kyphosis and lordosis-also contribute
the spring-like capacity of the spine and allow th
vertebral column to withstand higher loads than if
were straight. A study of the capaci.ty of cadav
thoracolumbar spines devoid of muscles to resi
vertical loads showed lhat the critical load (th
point at which buckling OCCUlTed) was approx
',:
,.
FIG. 10-13
The line of gravity for the trunk (solid line) is usually ventral
to the transverse axis of motion of the spine and thus the
spine is subjected to a constant forward-bending moment.
cles, the abdominal muscles an.: often interm
,-lclive in maintaining tlte neutral upright p
and stabilizing the trunk (Cholc\\'icki el aI.,
However, this activit~· is readily reduced by th
mand to stand relaxed (Hodges &. Richardson
The venchr;:ll portion of the psoas muscles
invol\'ed ill producing postul'al s\\'a~' (Bas
1958; Nachemsoll, 1966). The level of aCl
these muscles varies considerably among in
als and depends to some extent on the shape
spine, for eX<:\J11plc, on the magnilllde of h
kyphosis and lordosis.
The pelvis nlso plays a role in the muscle
and resulting loads on the ~pine during s
(Fig. 10-14). The base of the s<lcrum is inclin
ward and downward. The angle of inclinat
sacral angle, is approximatcl~1 30° to the tr~
plane dudng relaxed standing (Fig. 10-148).
or the pelvis abollt the tnmSVL:rse axis betw
hip joints changes the angle. \Vhell the p
tilted backward, the sacr<ll angle decreases
lumbar lurdosis flallens (Fig. 10-14..1). This
ing afrects the thoracic spine. which extends
to <'Idjust the ccnter of gravity or th<.' trunk
the cncrg,\' expenditure, in tcrms or muscle e
/'
.......
)i
~ :
((
\ ;,;
!
\ i
'J
.~
A
LOADING OF THE SPINE DURING STANDING
When a person stands. the postural Illuscles are
constantly active. This activity is minimized when
the body segments arc well aligned. During stand·
ing, the line of gravity of the trunk usually passes
ventral to the center of the fourth lumbar vertebral
body (Asmussen & Klausen, 1962). Thus, it falls
ventral to the transverse axis of motion of the
spine and the motion segments are subjected to a
forward-bending moment, which must be counterbalanced by ligament forces and erector spinae
muscle forces (Fig. I 0-13). Any displacement of the
line of gravity alters the magnitude and direction of
the moment on the spine. For the body to return to
equilibrium, the moment must be counteracted by
increased muscle activity, which causes intet"mittent
postural sway. In addition to the erector spinae mWi·
~
r'"
!'. \~\,
mate!y 20 to 40 N (Gregersen & Lucas, 1967; Lucas
& Bresiel; 1961). The critical load is much higher in
vivo and varies greatly among individuals. The ex·
trinsic support provided by the trunk muscles helps
stabilize and modify the loads on the spine in both
dynamic and static situations.
B
j
o
-30
Effect of pelvic tilting on the indination of the ba
sacrum to the transverse plane (sacral angle) durin
right standing. A, Tilting the pelvis backward redu
sacral angle and flattens the lumbar spine. D, Duri
laxed standing, the sacral angle is approximately 3
Tilting the pelvis forward increases the sacral angl
centuates the lumbar lordosis.
•
minimized. \Vhcn the pelvis is tilted ronvard, the
angle increases, accentuating the lumbar 101'and the thoracic kyphosis (Fig. 10-14C). Forand backward tilting of the pelvis influences
activit)! of the postural muscles by affecting the
loads on the spine (Floyd & Silvel; 1955).
Values of Intradiscal Pressure for Differe
Positions and Exercises As a Percentage
Relative to Relaxed Standing in One
Subject (Chosen Arbitrarily As 100'1',)
i Position/Maneuver
vu~""u
LOADS ON THE LUMBAR SPINE
STANDING, SIDING, AND RECLINING
position affects the magnitude of the loads on
spine. As a result of in vivo intradiscal pressure
studies conducted by! Nachemson
1975), it was shown that these loads are minimal
well-supported reclining, remain low during
re"",";u upright standing, and rise during sitting. A
recent in vivo investigation of intervertebral disc
pressure utilizing more sophisticated technology,
based on only one subject, suggested that in
relaxed, unsupported sitting, interdiscal pressure
is less than in standing (Wilke et aI., 1999). Aclditional pertinent pressure measurements can be
seen in Table 10-1. Sato et al. (1999) have verified
Nachemson's (1975) findings, showing an increase
in spinal load from 800 N in upright standing to 996
N in upright siuing. The relative loads on the spine
during various body postures, as described by
Nachemson and \Vilke, are presented in Figure 10-15.
During relaxed upright standing, the load on the
third and fourth lumbar disc is almost twice the
weight of the body above the measured level
(Nachemson & Elfstrom, 1970; Nachemson & MOiTis,
1964; Wilke et aI., 1999). Trunk flexion increases the
load ancl the fOl"\vard-bending moment on the spine.
During forward flexion, the annulus bulges ventrally
(Klein et aI., 1983) and the central portion of the
disc moves posteriorly (Krag et aI., 1987) (Fig. 10-16).
iVlore than trunk extension, trunk flexion stresses
the posterolateral area of the annulus fibrosus. The
addition of twisting motion and accompanying
torsional loads further increases the stresses on
the disc (Andersson et al.,1977; Shirazi-Adl. 1994;
Steffen et aI., 1998) (Case Study 10-2).
The loads on the lumbar spine are lower more during supported sitting than during unsupported sitting. During supported sitting, the weight of the upper body is supported by the backrest, which reduces
the muscle activity, relieving intradiscal pressure
(Andersson et aI., 1974; Wilke et aI., 1999). Backward
inclination of the backrest and the use of a lumbar
support further reduce the loads. The use of a support in the thoracic region, ho\vevel~ pushes the thoracic spine and the trunk fOl\vard and n1akes the
lumbar spine move toward kyphosis to remain in
lying supine
Percent
------20
Side-lYing
lying prone
lying prone. extended back,
supporting elbows
laughing heartily, lying laterally
Sneezing, lying laterally
Peaks by turning around
24
22
50
30
76
140-1
Relaxed standing
100
Standing, performing Valsalva maneuver
184
Standing, bent fOIVvard
220
Sitting relaxed. without back rest
92
110
166
86
Sitting actively straightening the back
Sitting with maximum flexion
Sitting bent forNard 'vvith thigh
supporting the elbows
Sitting slouched into the chair
Standing up from the chair
Walking barefoot
Walking with tennis shoes
Jogging with hard street shoes
Jogging with tennis shoes
Climbing stairs, one at a time
Climbing stairs, two at a time
Walking down stairs, one at a time
Walking down stairs, two at a time
lifting 20 kg, bent over with round back
lifting 20 kg as taught in back school
20 kg close to the body
Holding 20 kg, 60 em away from the chest
Holding
Pressure increase during the night rest
(over a period of 7 hours)
106-1
106--1
70-19
70-17
100-1
60-24
76-12
60-18
460
340
220
360
20-48
Adapted '.vith permission from Wilke, H.J .. Nee!. P. Caimi, M.. et
(1999). Ne\lv in vivo measurements of pressures in the interverteb
disc in daily life. Spine, 2";, 755.
-----------_._-
~
500
0 Nachemson, 1975
"(;9.-
400
0
ro"
,~
'1"
i~
]
200
E
0
z
,~
"
1
100
A
~
~~~~
.~
~
~,
%
',Z
;) ~
;,
"
~
c
~
#~
ter working
a 12-hour shift, when he lifted \-vhile twistin
~
an unusually large, yet lightweight box, During the iirst
J.!
week of pain, he visited his physician, who prescribed pa
c!ll
,;I
,.i
•
35-year-old male presents complaints of low back
pain with radiation lO the posterior aspect of the le
thigh, not past the knee. His pain staned 3 weeks ago,
%
300
'"
Nonspecific Low Back Pain
~
Wilke, 1999
'"
iii
£
'C
.
".~
,S
en
c
'5
c
~-----'-----'-'-'--
~
,
'1
medication. Currently, he is still in pain, panieufarly durin
sitting
or standing for loog periods. During a iollmv-up
physician visi!,
a careful examination showed
the patien
to be someV'ihat ovel\veight. ",vith weakness in his abdo
~
ina I and back muscles and poor flexibility in hiS ham'1
strings. psoas, and back muscles. Neurological tests were
normal and diagnostic x-rays were normal as well, leadin
to a diagnosis of nonspecific low back pain (Case Study
Fig.10-Z-1).
Data from two studies using intradiscal pressure measurements. The relative loads on the third and fourth lumbar
discs measured in vivo in various body positions are compared with the load during upright standing. depicted as
100%. Adapred wirh permission {rom Nachemson. A. (1975).
TOWMds a berrer unrJersfc1nding of back pain: A review of rile
mechanics of the lumbiJr disc. RheumaroJ Rehabil, 14, /29 and
{rom Wilke, H.J.. Nee!. P.. (aimi. M., ei al. (999). Ne~·'/ in vivo
medSuremen!s 0; preSSlJreS in rhe inrerve(rebral disc in daily life.
Spine, 24, 755.
•
)
Tension
Compression
, Case Study Figure 10-2·1.
Combinations of different factors have resulted in th
injury. From the biomechanical point of view, although
the load lifted was considered light, the vastness of the
The forward-bent position produces a bending moment on
the lumbar spine. The moment is a product of the force
produced by the weight of the upper body (W) and the
lever arm of the force (lw). The forward inclination of the
upper body subjects the disc to increased tensile and compressive stresses. The annulus bulges on the compressive
side and the nucleus is shifted posteriorly.
package and the resultant large lever arm (the distance
from the center of gravity of the person to the package
crea'ted a larger than expected load on the lumbar spine
In addition, weakness in the abdominal and extensor
muscles of the spine led to an additional mechanical dis
advantage in stabilizing the lower back. Tight psoas and
hamstring muscles place restrictions on the mObility of
pelvis, stressing the range of motion jn the lumbar regio
and affecting the normal loads and motions at this leve
·i·
.,.
-~
.-;
conwct with the backrest. increasing the loads on the
lumbar spine (Andersson et aI., 1974) (Fig. 10-17).
Loads on the spine arc minimized when an incliassumes a supine position because the loads
b)' the bod;!s weight arc eliminated (Fig,
With the body supine and the knees ex>tenden, the pull of the vertebral ponion of the psoas
muscle produces some loads on the lumbar spine.
the hips and knees bent and supported, howthe !umb4\r lordosis straightens out as the psoas
>'n'luscle relaxes and the loads decrease (Fig. 10-18).
!
'"-~--.~.
I
----,.
B
FiG."'10.18
STATIC LOADS ON THE LUMBAR SPINE
DURING LIFTING
The highest loads on the spine are generally pro!t~>;.,S7 duced by c:'\ternalloads, such as lifting a heavy object. Just how mllch load can be sustained by the
spine before damage occurs continues to be investigated. Pioneering swdies by Eie (1966) of lumbar
ii
~
fil
II'I
III
,1J
i
/j:
!i))i
r~1
'-'...J j
! '
A
B
I~Di~sc1P;re~ss~u~re=========~;I:'===1~
~
;,
:
Influence of backrest inclination and back support on
loads of the lumbar spine. in terms of pressure in the third
lumbar disc, during supported sitting. A, Backrest inclination is 90" and disc pressure is at a maximum. e, Addition
of a lumbar support decreases the disc pressure. C, Backward inclination of the backrest is 110", but with no lumbar support it produces less disc pressure. D. Addition of a
lumbar support with this degree of backrest inclination
further decreases the pressure. E, Shifting the support to
the thoracic region pushes the upper body forward, moving the lumbar spine toward kyphosis and increasing the
disc pressure. Adapted with permission from Andersson, G.8.1..
Ortengren, R.. Nc1CiJemson, A.. et at. (1974). Lumbar disc pressure and myoelectric back muscle activity during sitting. 1. SHldies on cln experimental chair. Scand J Rehabil Med, 6. 104.
A, When a person assumes a supine position with legs
straight, the pull of the vertebral portion of the psoas
muscle produces some loads on the lumbar spine. B, Whe
the hips and knees are bent and supported, the psoas mu
cle relaxes and the loads on the lumbar spine decrease.
I
•
vertebral specimens from adult humans showe
that the compressive load to vertebrae failur
ranged from approximately 5,000 to 8,000 N. O
the whole, values reponed subsequently by othe
authors correspond to those of Eie, although va
ues above 10,000 N and below 5,000 N have bee
documented (Hullon & Adams, 1982). The appl
cation of static bending-shearing moment on lum
bar motion segments revealed that bending mo
ment of 620 Nm and shear rnoment of 156 Nm
were tolerated before conlplete disruption of mo
tion segmel1l occurred. The flexion angle befor
failure was recorded as 20° with 9 mm of horizon
tal displacement bct\veen the two vertebrae (0$
valder et aL, 1990). 80th age and degree of disc de
generation innuence the range preceding failure
Although the vertebral body strength is relative t
Ihe bone mass, with aging ~ the decline in bon
strength is more pronounced than is the decline i
bone mass (iVlosekildc, 1993).
Eie (1966) and Ranu (1990) observed that durin
compressive testing the fracture point was reache
in the vertebral body, or end plate, before the inte
vertebral disc sustained damage. This finding show
that the bone is less capable of resisting compres
sian than is an intact e1isc, During the testing, a yiel
point was reached before the vertebra or end pla
fractured. \Vhen the load was removed at this poin
the vertebral body recovered but was morc suscep
lible to damage when reloaded.
Evidence exists that the spine may incur microdamage as a result of high loads in vivo. T. I-Jansson
(1977, The hone mineral contenl (Int! biomec!tanica!
properties
lumbar vertebrae. All ill vitro study
based on dllal p1W101/ absorptiometry. Unpublished
thesis, University of Gothenburg, Sweden) observed
microfractures in specimens from "noI'mal" human
lumbar vertebrae and interpreted this microdamage
to be fatigue fractures resulting from stresses and
strains on the spine in vivo. In vitro examination
confirmed the cxistance of microdamage near the
end plated with compression loading (Hasegawa et
aI., 1993).
Lifting and carr.ying an object over a horizontal
distance arc common situations wherein loads applied to the vertebral column may be so high as to
damage the spine. Several factors influence the
loads on the spine during these activities: (t) the position of the object relative to the center of Illotion
in thc spine; (2) the sizc, shape, weight. and density
of the object; (3) the degree of flexion or rotation of
the spine; and (4) the rate of loading.
Holding the object close to the body instead of
away from it reduces the bending moment on the
lumbar spine because the distance from the center
of gravity of the object to the center of motion in the
spine (the lever arm) is minimized. The shorter the
lever arm is for the force produced by the weight of
a given object. the lower the magnitude of the bending moment and thus the lower the loads on the
lumbar spine (Andersson et al., 1976; Nachemson &
Elfstriim, 1970; Nemeth, 1984; Wilke et aI., 1999)
(Calculat ion Box 10-1).
Even when identical and nonfatiguing repeated
lifting tasks are perFormed, variability in lifting technique of the same subject has been shown in trunk
kinematics, kinetics, and spinal load (Granata ct aL,
1999). ,",Vhen an individual repeatedly performs an
identical lift, great variability is recorded, which indicates that the brain may have several motor strategies to do a task. It also indicates the sensitive responsiveness of the muscle system to subtle changes
to maintain the performance despite fatigue.
,",Vhen a person holding an object bends forward,
the force produced by the weight of the object plus
that produced by the weight of the upper body create a bending moment on the disc, increasing the
loads on the spine. This bending moment is greater
than that produced when the person stands erect
while holding the object (Calculation Box 10-2).
Health professionals generally recommend that
lifting be done with the knees bent and the back relatively straight to reduce the loads on the spinc.
or
Influence of the Size of the Object
on the Loads on the Lumbar Spine
The size of the object held influences the loads on the lum-
bar spine. If objects of the same weight, shape, and density
but of different sizes are held, the lever arm for the force
produced by the weight of the object is longer for the
larger object, and thus the bending moment on the lumba
spine is greater (Calculation Box Fig. 10-1-1). In these t\VO
situations (Calculation Box Figs. 10-1-1 ,md 10-1-2). the
distance from the center of motion in the disc to the front
of the abdomen is 20 em. In both cases, the object has a
uniform density and weighs 20 kg. In the case of Calculation Box Figure 10-1-1, the \·vidth of the cubic object is 20
crn; in the case of Calculation Box Figure 10-1-2, the width
is 40 ern. Thus, in Case 1 (Calculation Box Fig. 10-1-1), the
forv·.wd-bendlng moment ,Kung on the 10'/'Jest lumbar disc
is 60 Nrn, as the force of 200 N produced by the weight o
the object acts \vith a lever arm (U of 30 cm (200 N
;<
0.3
m).ln Case 2 (Calculation Box Fig. 10-1-2). the forvvardbending moment is 80 Nrn, as the lever arm (L;.) is 40 ern
(200
(\J
x 0.4 rn).IConsidering 1Kg
10N.]
20 em
200 N
Forward-bending
moment 60 Nm
Calculation Box
Figure 10-1-1.
Forward-bending
moment
80 Nm
Calculation Box
Figure 10-1-2.
I
~ :./.;~; :,
!ji/! I ~:I~~;Cto~::haet~h~e:u~::r~;i~~on
lii}i!, I
;:~ n~ ~O~; ,:;: dSehn~: ~1 i:b~:~~::;h~nBgO;~i~~r~s
10-
~:
lifted, In Case 1 (Calculation Box fig. 10-2-1) (upright
f
I
standing), the lever arm of the force produced by the.- f
i' '/
bending moment of 60 Nm (200 N
weight of the objecr (lp) is 30 em. creating a forward-
r
x
0.3 m), The
forward-bending moment created by the upper body is
~',
9 Nm; the length of the lever arm (LW) is estimated to
t',;
body is 450 N. Thu5, the tolal forward-bending moment
•
in Case 1 is equal to 69 Nm (60 Nm + 9 N(m).
b d
In Case 2 (Calculation Box fig. 1O-2 -2J upper 0 y
~
2 em, and the force produced by the weight of the upper
J
~.'
flexed forward), the lever arm of the force produced by
f
I
the weight of the object (l..) is increased to 40 em, creat-
Ii
ing a fon,.vard·bending moment of 80 Nm (200 N x 0.4 m).
furthermore, ,he force of 450 N produced by the weight
of the upper body increases in importance as it acts with
r
a lever ann (L...) of 25 em, creating a forward-bending
moment of 112.5 Nm (450 N x 0.25 m), Thus, the total
I
ff,
~:
forward-bending moment in Case 2 is 192.5 Nm (1125
~'M""
t
I
I
"
However, this recommcndalion is valid only if this
technique is used correctly and optimally, with the
load positioncd betwcen the feet and thcrcby reducinn the lever arm of the extcrnal load (van Dieen ct
al~ 1999) (Calculafion Box 10-3),
A literature review revealed no significant differcnce in spinal compression and shear computed
forces betwcen SlOOp or squat lifting (van Dicenc ct
at., 1999). However, it was suggestcd lhat loss or balance is more likely during squat lifting. which in
turn m,,)' ,-,dd addilional stresses on the lumbar
spine.
In thc following examplc, thc frcc-body technique for coplanar forces will be used to make a
simplified calculation of the static loads on the
spine as an object is lifted (Calculation Boxes 10-4;
and 10-4B).
Calculations made in this way for one point in
lime during lifting arc valuable for demonslrating
how the lever arms of the forces produccd by the
wcight of the uppcr body and by the weight of the
obj;ct affect the loads i~,posed on the spine, The
usc or the salllc calculations to compule the loads
produced when an 80 kg objcct is lifted (representing a force of 800 N) yields an approximate load o
10,000 N on the dise, which is likely to exceed thc
fracture point of the vertebra. Because athletes who
lift wcights can easilv reach such calculated loads
without sustaining fr~ctures, other factors, such as
intra-abdominal p~-essurc (lAP), may be .involved in
reducinrr thc loads on the spine in vivo (Krajcarsk
et aI., 1999),
Dynamics
~Lw
<:r'=i~-A-=-::t=-=-i
J
450 N
Almost all mol ion in lhe body increases muscle rc
cl'uitmcnt and the loads on the spine. This increase
is modest during such aclivities as slow walking
or casy twisting but becomes morc marked during
variou's exercises and the coqlplexity or dynamic
movement and dynamic loading (Nachclllson &
Elfstrom, 1970),
WALKING
,I
200 N
I
I
i
Tolallorwardbending moment
= 192,5 Nm
Talal fo"vardbending moment
= 69 Nm
11
Calculation Box
Figure 10-2-1.
Calculation Box
figure 10-2-2.
$---------"-~---_.
In a study of normal walking at four speeds, the
compressive loads at the L3-L4 mOlion segmen
rangcd from 0.2 to 2.5 limes body weight (Fig
10-19) (Cappozzo, 1984), Thc loads were maxima
around loe-off and increased approximatel~1 lin
early with walking speed. Muscle .action was
mainly concentrated in the trunk extensors. Incli
vidual walking trailS. parlicularly the amount o
forward flcxion of the trunk, influenced thc loads
The Technique Employed During Lifting Influences the Loads
on the Lumbar Spine
In the three situations shown in Calculation Box Figures
total forward· bending moment of 151 Nm(f200 N
10-3-1, 10-3-2. and 10-3-3, an identical object weighing
20 kg is lifted. Case 1 (upper body flexed forward) (Caleula-
+ 1450 N x 0.18 m!).
Case 3 (Cal(l,lIalion Box Fig. 10-3-3) shows that be
tion Box Fig. 10-3-1) is identical to Case 2 in Calculation Box
knees per se do not decrease the forward-bending mo
x
10-2, v~here ttle total fONJarcl-bending moment is 192.5 Nm.
If the object lifted is held out in front of the knees, the
In Case 2 (Calculation Box Fig. 10- 3-2), lifting with the
arm of the force produced by the weight of the objec
knees bent and the back straight places the object closer to
creases to 50
the trunk, decreasing the forwclrd-bending moments. The
the weight of the upper body (L..:) increases to 25 em.
ern.
and [he lever arm of the ierce produ
lever arms of the forces produced by the weight of the object
tile total forwarcj-bending moment created is 212.5 N
(U and Lipper body IL,) are shortened 10 35 and 18 em, re-
1{200 N x 0.5 mj ... 1450 N x 0.25m!).
spectively. at this point in the lifting process. The result is a
.';'
200 N
200 N
Total
forward·bending
moment = 192.5 Nm
Total
forward-bending
moment = 151 Nm
Total
forward-bending
moment =: 212.5 Nm
Calculation Box
Figure 10-3·1.
Calculation Box
Figure 10-3-2.
Calculation Box
Figure 10-3-3.
0'-~~~~~~~~~~~-
The greater this flexion, the larger the muscle
forces and hcncc the compressive load. Callaghan
et al. (1999) corroborated lhese findings and further showed that walking cadence affects lumbar
loading, with increased anterior-posterior shear
forces noted as speed increased. Limiting arm
swing during walking resulled in increased com-
,-
pressive jomt loading and EMG outpu
creased lumbar spine motions. 1n conc
cause of low tissue loaeling, walking is
perhaps ideal therapeutic exercise ror
low back pain (Callaghan et al" 1999) w
tion to speed of walking can further
spinal loads (Cheng ct al" 1998)_
Diagram Technique for Coplanar Forces. Calculation
the Static Loads on the Spine As an Object Is Lifted
loads on cl lumbar disc I,'iill be calculilled for on,! point in tim[' 't/hen a
NUJ to be posit]>,'e and counterciochvis(' moments are considered to
cer50n who weighs 70 kg lifts a 20 kg obj!'ct The spin£' is flexed ilpprW.i.
negeltl'/e.i
,
Thus,
35 In this example, the three priflclpdl forces ,Kling on the lurnl)ilf
Q
allhe lurnbosilual le'lei ilfe: (1) HK force produced by \1',(' ','ieight 01
upper body (Wi. Ciilculat.:d to be ':50 N (approxin\ille!y 65",0 of the
exerted by lhe IOtal body weigh!); (2) tile fOlce produced b~i the
of the object (Pl, 200 N; "neJ (3)
trw
force produced by contraction
the erector spinae muscles (El, which has a kno':m direction anel point of
0
L)
\t ;.: 0.05 m)
(450 N ;.; 0,25 m) " (200 N >: DAm)
E ;.: 0,05 rn
112.5 Nrn -;- 80
0
1~m
Solving this equation for E yields 3.850 N
ap,al",,'''''' but an unknQl,W1 magnitude (Calculi1lion Box Fig. 1O·4A·l)
a distance
0
:::: r,il
{\,V :< L<..J· (P >:U - iE >.:
from the center of motion
Th!: to",1 compre~~!'/e force exe!led on the disc (Cl can now be calcu-
in the spine, they crcate moments in the lurntBr spille, Two fOPNard-ber:(j·
lawd trigonof1,etricaliy (Calculation Box Fig. 10·cIA-2). in the example, C is
in,) moments (\iVl:,_ cl!ld PL,) are the products of (V{: and (Pi and the perpen-
lhe sum of lhe compressive forCi':s "cLng O'lel the disc. '.vhich is inclined
diculars from the instanl center of rOIMion to the lines of action of these
to the triirlSVelSe pldfle These forces are
Because these three forces <let at
fc,rces (,heir lever
Mms),
The le'/er arm (L'J for (Pl is 0.<1 III and the !e'/er arm
The compressive force produced by the weight of the upper body
(L,,) for (W) is 0.25 m. f\ counterbalanCing moment (ELi is the product of
(W:-: cos 35°).
2 The force producHJ by the '.veight of the object (Pl, which acts on
the disc inclined 35" (P '< cos 35"')
(V'/), \'ihich acts on the disc inclined 35'0
lEi and Its lever arm, The lever elm: (l) is 0,05 m, TIl(' fI1agnitude of (E) CiJll
b~~ found tilrough the use of the equilibriufI1 equa,ion for moments, For
,he
kJdy to be in moment equilibrium, the sum of the moments aCiing on the
;u:nbar spine must be
zeo
Th;;,: fOlce produced by ,he erector spinae muscles (E), 'i,hich acts ap-
3
(In ihis example. cloch'!ise morm,nts "re consid-
proxirn,w.:ly at il right angle to the diSC inclination.
The totdi compressive fOlce acting on the disc (C) has a known sense,
point of ilppliciltion, dnd line of iKlion but an unknown magnitude. The
n:dgnllude of C (an bt": ioundlflrough H\e use of the eq\Jiiibriurn 0quation
for forces For the body to be in force eqUilibrium, the sum of the forces
must b(, .-,qual to zero
Force W (450 N)
Thus,
a
:::: forces
f
(v') :-: cos
Force C
(magnitUde
unknown)
I \,
L-!
\,
\
\
l'
C '" 4382 N
S
373 N
\
~P
Force P (200 N)
(';501'-1 ;.: cos
~
C
Force S
(magnitude
unknown)
3S~) .c.
cos 35") .. E
C
°
(200 N ;.: cos 35") -'. 3850 N ." C
;co
0
368.5 N .;. 163,8 N·· 3850 1'J
Solving the equation for C yields 'U82 1'1
The shear component for the reaction force on the disc (S) is found in the
same'i/ay
{450
Calculation Box
Figure 10-4A-2.
Calculation Box
Figure 10·4A~1.
35") ,. (P:·
~J
'.,: sin 35") ;, {ZOO N :': sin 35") _. S ,'.
°
373 N
~r-----------------------------------------
Free~Body Diagram Technique for Coplanar Forces. Calculation
of the Static Loads on the Spine As an Object Is Lifted
8ecause C and S form" right angle (Calculation 'Box Fig 10·48·1). the
Forco S
j
85
35
35
Forco W • Force P
/
/
vc .;. s;'
{R'I
R
. ; ; Forco R
Forco<;!/ .
Pythagorean theorem can be used !O fin,J the towl reaction force on the disc (Rt
Forco
4398N
The direction of (R) is determined by means of a trigonometric function:
/
sin h) '" CJR
FOTeo E
(tt) '" archsin (ClR)
;co
85" where (t,) is the angle belween lhe total vecto
force on the disc and the disc inc!ina,ion
The problem can be graphically solved by COlf5trUCling a v0ctor diagram
b"sed on the knQ'.'m values (Calcul,llion Box Fig, 10-48·2), A venicalline repre
35
Calculation Box
Figure 10-48-1.
senting (V'l) ,-;. (P) is drawn first;
Calculation Box
Figure 10~4B·2.
(EJ
is added at a right angle to the disc inclina
tion, and (R) closes ,he triangle. The direction of (R) in relation to the disc
termined
Walking Speed
Slow (1.05 mfsec 1)
Normal (1.38 m/sec' l )
--------Fast (1.72 m/sec' l )
Very fast (2.16 mfsec 1)
:c0>
·ffi
2
;;:
>-
"In0
<5
~
E
~
f-
.£
~
~
0
LL
0
RHS
LHS
LHS
~0l-L---:':--'-l.:-I-:'::-'-:!:::-''---:JL
20
40
60
80
100
Percentage of Cycle
Axial load on the
L3~L4
motion segment in terms of body
weight for one subject during walking at four speeds. The
horizontal line (UBW) denotes the weight of the upper
body, which represents the gravitational component of
this load. Loads were predicted using experimental data
from photogrammetric measurements along with a biomechanical model of the trunk. LHS, left heel strike; RHS,
right heel strike. Adapted with permission from Cappozzo, A.
(1984), Compressive toads in the lumbar vertebral column during normal level walking. J Orthop Res, 1, 292.
EXERCISES
During strengthening exercises for the erector
spinae and abdominal muscles, the loads on the
spine can be high. Although such exercises must be
effective for strengthening the trunk muscles con~
cerned, they should be performed in such a wa:\' that
the loads on the spine are adjusted to suit the condition of the individual.
The erector spinae muscles are intensely activated
by arching the back in the prone position (Fig.
IO-20A) (Pauly, 1966). Loading the spine in extreme
positions such as this one produces high stresses on
spine structures, in particular the spinous process
(Adams et aI., 1988). Although intradiscal pressure
in a prone position with upper bod)' support on the
elbows is halF that in standing (Wilke et aI., 1999), it
is recolllmended for c:\ercise that an initial posi
that h:eps the \"e["IL'hr~IL' in a 1l1Orc parallL'1 alignn
is preferahle \"hen stn.:ngthcning e:\L'rcises for
LTector spinae IllLlscles ~\re performed (Fig. 10-2(
The importance of the ahdorninal muscles
spinal stahilit.'· and inlL'rpla.\" in the production
lAP reinforces the need ror strong ahdorninal l
ors. Sit-ups arc a userul exercisc ror ahdcllll
rnusclc strengthening, wilh man." vari:'\lions p
ticed and cncOllraged by health prol"cssionals,
cerlain vari:'l1ions arc viewed as hannflll to the
back. Although the most common belief has h
that sit-ups with the knees llexed and fed ll
chored will emphasize the abdominal contribut
while minirnizing psoas acti\"it." (.JukeI' et aI., 19
this is not true. Both bent knee and straight leg
ups will producL' comparable le\"L'ls of PSO~\s and
domin~d aClivit.\", crealing c(lmpressivc spinal lo
ing. Curl-ups, in which the head and shoulders
raised onl.\" to the point where the shoulder bla
clear the l<.lble and lumb<.lr spinL' motion is m
mi!.ed (F'ig. 10·21) and ortell emphasized in rehab
tation programs, arc recommended for minimiz
compressi\"L' lumbar loading (A,der &. l\-!cGil!. 1
Jukcr L't aI., I (98). This modilic<'ltion of the exen
has been shown to be cl"fecti\'c in terms or Illotor
recruitment in the muscles (F.kllOlm el aI., 1
Flint, 1965; Partridge &. \\'alters, 1959); all port
01' the external oblique and rcctus abdolninis rnus
<.lre activated. Sit-ups with feet unanchored, legs
vated, or tOI"SO twisting do not significantl.v incre
abdorninal muscle <.lclivit." (AxleI' &: J'dcGill, 1997)
lirnit the psoas activity, <.l reverse curl, wherein
knees arc brought t()\vard the chest and the butto
<.lre raised I"rom the table, aetivalL's the internal
external oblique rnuscles :.md the reclus abdom
muscle (P<.\rtridge &: \Vahers, 1959), If the rev
curl is perrormcd isometricall.v. the disc pressur
lower than that produced during a sit-up, bUl the
ercise is just as effective for strengthening the
dominal rnuscles (Fig. IO~22l. It can he conclu
that no single :.Ilxlominal exercise C<.m optimall.v t
~dl tnmk flexors while minimizing intervertebral j
loading. Instead, a \'aried program must he
scribed, tailored to the training objectives of the
dividual (;hler & McGill, 1997).
\Vhen designing a back strengthening cxer
progr:'llll, the most import:.lnt consider:.ltion is
conclusion drmvn bv the Paris Task Fo
(Abenh:.lim et al., 2(00). The guidelines set forth
clude recommendations that exercise is benefi
for subacute and chronic low Inlck pain, No par
ular group or t~'pe of exercises has bcen shown t
most efTccti\·c.
A
A, Arching the back in the prone position greatly activates the erector spinae muscles but
also produces high stresses on the lumbar discs, which are loaded in an extreme position.
B, Decreasing the arch of the back by placing a pillow under the abdomen allows the discs
to better resist stresses because the vertebrae are aligned with each other. Isometric exercise in this position is preferable.
.>
,?
.',.
200
150
100
50
Performing a curl to the point where only the shoulder
blades dear the table minimizes the lumbar motion and
hence the load on the lumbar spine is less than when a
full sit-up is performed. A greater moment is produced if
the arms are raised above the head or the hands are
clasped behind the neck, as the center of gravity of the
upper body then shifts farther away from the center of
motion in the spine.
A reverse curl, isometrically performed, provides efficient
training of the abdominal muscles and produces moderat
stresses on the lumbar discs. The relative loads on the thir
lumbar disc during a full sit-up and an isometric curl are
compared with the load during upright standing, depicted
as 100%. Adapted with permission from Nachemson, A. (1975
Towards a better understanding of back pain: ~ review of the
mechanics of the lumbar disc. Rheumatol Rehabil, 14. 129.
•
........
...
. ".~-
ey,.
MECHANICAL STABILITY
OF THE LUMBAR SPINE
i\tlechanical stability for the lumbar spine can be
achieved through several means: lAP, co·conlraclion of the tnlllk muscles, external support, and
surgery. Surgical procedures for lumbar spine stability will not be covered in this section.
Intra-Abdominal Pressure
lAP is one mechanism that may contribute to both
unloading and stabilization of the lumbar spine.
.IAP is the pressure created within the abdominal
cavity by a coordinated contraction of the diaphragm and the abdominal and pelvic floor mus-
cles. [lS unloading mechanism was first propos
Banelink in 1957 and I\-torris et al. in 1961.
suggested that lAP serves as a "pressurized ba
altempling lO separate the diaphragm and
noor (Fig, 10-23, A & B), This creates an ext
moment that decreases lhe compl-essio!1 forc
lhe lumbar discs_ The extensor moment pro
by TAP has been calculaLCd in several biomecha
models, with widely varying resulting reductio
extensor moment from 10 to 400;' of the ex
load (Anderson et at .. 1985; Chaffin, 1969; Eic,
Lander et al .. 1986; Morris et aI" 1961).
Recent studies using fine-wire EMG o
deeper abdominal muscles found that the tran
sus abdominis is the primary abdominal musc
sponsible for lAP generation (Cresswell,
Spine
160
140
Force
."
-lbs- --::._.
.......
120
"
100
80
Abdominal
cavity
60
40
Respiratory flow
Positive
values = Inspiralion
Negalive
values=expiralion
20 r
Line of lorce
0
\r-
n:--~ Inlraabdominal
pressure
: -mmHg-
!
",,'
lime
-msec-
Ibs
mmHg
A
B
A, Schematic illustration of the effect of intra-abdominal pressure. An increased pressure will create an extension moment
on the lumbar spine. B, Intra-abdominal pressure (lAP) (measured by a nasogastric microtip transducer) and respiratory
flow (measured through a Pneumotach) during stoop lift of
120 lbs. (approximately 60 kg). Solid line, lAP; dotted line,
force exerted in lbs; dashed line, respiratory flow (negative
values delineate expiration and positive values delineat
ration). Note that the subject inspires before the lift an
the breath throughout the lift. lAP increases and peaks
gether with the lifting force. helping to stabilize and un
the lumbar spine. (ourtesy of Markus Pieuek, MD and Mar
Hagins PI: MA. Program of Ergonomics and Biomechanics, N
York. UniversiC'j and Hospital for Joint Diseases, New York, Ny
Cresswell el aI., 1994a; Cresswell eL aI., 1992;
Hodges cl aI., 1999). As the transversus abdominis
is ho.-izontally oriented. it creates compression and
an incre,lSC in lAP withollt an accompanying flexor
moment. A recent experimental study gave the.firsl
direct evidence ror a trunk extensor moment produced by eleva Led lAP (Hodges el aI., 2000),
It has been demonstrated that the lAP contributes LO the mechanical stability of the spine
through a co-activation between the antagonistic
lrunk lIexor and extensor muscles (Cholewicki et aI.,
1997, 1999a,b; Gardner-Morse & Stokes, 1998). As
the abdominal musculature contracts, JAP increases
and converts the abdornen into a rigid cylinder that
greatly increases stability as compared with the
muhisegmenlcd spinal column (ivlcGiII & Norman.
1987; Morris et al.. 1961).
lAP increases during both stalic and dynamic
conditions such as lifting and lowering, running Hncl
jumping, and unexpected trunk perturbations
(Cresswell el aI., 1992; Cresswell et aI., 1994b; Cresswell & Thorslcnsson, 1994; Harman et aI., 1988).
Current research suggests that the transversus ab~
clominis muscle, together with the diaphragm, plays
an important role in stabilizing the spine in preparalion for limb movement, regardless of the direction in which n1QVement is anticipated. Transversus
abdominis and diaphragmatic activity appear to occur independently, prior to activity of the primary
limb mover or the other abdominal muscles
(Hodges eL al.. 1997, 1999).
Trunk Muscle Co-Contraction
To understand the phenomenon or co-contraction
during trunk loading, Krajcarski et al. (1999) studied the in vivo muscular response to perturbations
at two rates causing a rapid flexion moment. The
results of maximum trunk flexion angles and resulting extensor moments were compared. The results showed that with higher levels, co-contraction,
spine compression. and trunk muscle stiffness increase. During unexpected loading, a 70% increase in muscle activity has been noted as compared with anticipated loading, which may lead
to injury (Marras et al.. 1987). Further investigation into the loading response has revealed that
an inverse relationship (i.e .. the shorter the warning time, the greater the peak trunk muscle response) exists between peak muscle response and
warning time prior to loading (Lavender ct a\.,
1989).
Loss of spine stabilit~. . can be achieved throug
repetitive loading. This can be achieved throug
repetitive continuous motions that fatigue th
trunk muscles. l\tlusclc endurance is mechanicall
defined as the point at which fatigue of the mus
cles is observable. usually through a change i
movement pauern. Parnianpour ct al. (1988) use
an isoinenial triaxial device to study force outpu
and movement patterns when subjects performe
a flexion and extension rnovemcnt of the trunk un
til exhaustion. The results showed that with fa
tigue, coupled motion increased in the coronal an
transverse planes during the Oexion and extensio
movement. In addition, torque. angular excursion
and angular velocity of the motion decreased. Th
reduction in the functional capacity of the f1exion
extcnsion muscles was compensated for by sec
ondary Illuscle groups and led to an increascd cou
pIc motion pattern that is morc injury prone
Figure 10-24, A & B shows the increase in axial ro
tation (torque and position) during flexion and ex
tension of the trunk until exhaustion.
In an animal study, Solol11onow ct al. (1999) in
duced laxity of the spine in the ligaments, discs
and joint capsule by' cyclic repetitive loading of fc
line in vivo lumbar spines. The cyclic loading re
sulted in desensitization of the mechanoreceptor
with a significant decrease or complete elilnina
tion of reflexive stabilizing contractions of th
multifidus muscle. This may lead to increased in
stability of the spine and a lack of protective mus
cular activity even before muscular fatigue is ob
sen/cd. A IO-minute rest period restored the mus
cular activity to approximately 250/0.
External Stabilization
Restrict.ion of motion at any level of the spine ma
increase motion at another level. The usc of bac
belts as a means of preventing low back injury re
mains controversial. Originally it was believed t
assist in increasing lAP as a way of unloading th
spine during lifting; however, inconclusive ev
dence exists as to the biomechanical effectivenes
of these devices (Perkins & Bloswick, 1995). Th
National "Institute for Occupational Safety an
Health has advised against the use of back belts L
prevent low back injuries (NIOSH, 1994). As wel
an orthotic worn lo restrict lhoracic and lumba
motion may result in compensatory motion at th
lumbosacral level (Lumsden & Morris, 1968
Norton & Brown, 1957; Tuong et aI., 1998).
,
78.3
~
7.9
E
J\I\I~I~\~I\
!!
T
~
-·32.4
llmlm 1111111 Ill\/II\
!i~i I~i~
-'::======::;:======::;====================~
-127.1
58.6,
40.9
21.6
1/111/1111\
;111111
III I1II1
2.3
- 15.4
+--.,-'''''1-'--,'--'1-r'-'
, 'I'-'I--''--'--Ir-,TI--...,--,1+-i-.....-r-1T,- ,I-"--""I-,rl
0.1
9.3
18.4
27.5
36.6
45.7
54.7
A
ROlation Torque and Position vs. Time
4.3,.-------------------------------,
,
2.3
E
z,
"
~
1.1
!!
....0
-0.6
-2.1
19.1
,
w
13.3
'i'"
7.0
"co~
c
.Q
.~
0
Q.
0.7
- 5. 1 -!--,----,-.-,.--...,--,-.,..--,----,-.-,.--...,--,-.,..--,----,-'-.-,.--...,--,,-.,..--,----,--I
0.1
9.3
18.4
27.5
36.6
45.7
54.7
B
Dynamic (isoinertial) flexion-extension trunk testing until
exhaustion for one subject. Torque and position data is depicted for two planes, flexion-extension (A) and axial rotation (B). Note that flexion-extension torque production is
diminishing as is the amount of performed extension of
the trunk (A). Rotational torque and movement amplitude,
increased accessory motion, and torque is shown in B.
Data for lateral flexion was similar to axial rotation a
not shown here. Adapted with permission .from Parnianp
M .. Nordin. M.. Kahclnovitz, N.. er al. (1988). The rriaxial
coupling 0; torque generarion of trunk muscles during iso
ric exertions and the effect of fatiguing isoinercial movem
on the motor outpur and movemeor pc1tterns. Spine. 13(9
982-992
Investigation into the effect of back bells on
muscle activity has revealed no significant EMG
activity differences in the back extensors during
Iifling with or withoul a back bell (Ciriello &
Snook. 1995; Lee & Chen. 1999), while McGill el al.
(1990) showed slightly increased EMG aClivily in
the abdominal (except for the intel'l1al oblique
l11uscles) and erector spinae muscles. Thomas et '.11.
(1999) have verified a slight increase in EMG activity (2(10) in the erector spinae during symmetric
lifting wilh a back belt. Back bells have not been
shown to significantly increase Iifling capacity
(Reyna el aI., 1995).
on the spine to bc minimized during lifting, the d
tance bL'lween the trunk and thc object lifted shou
be as short as possible.
10 lAP and co-contraction of trunk musculatu
increases the slability of the spinal column.
1/1' Trunk muscle fatigue may expose the spine
increased vulnerability as a result of loss of' mot
control and thereby increased stress on the su
rounding ligaments, disc, and joint capsules.
,12 vValking is an excellent exercise thai poses
low load on the lurnbal' spine.
REFERENCES
Summary
1 The lumbar spine is a highly intricate and complex structure_
2 A vertebra-disc-vertebra unit constitutes a molion segment, the [unctionalunit of the spine.
, 3" The intervertebral disc scrves a hydrostatic
function in the motion segmcnt, storing energy and
distributing loads. This function is reduced with
disc dcgeneration.
~:"4 The primary function of the facet joint is to
guide the motion of the mOlion segment. The orientation of the facets determines the type of motion
possible al any level of the spine. The facels may
also sustain compressive loads, particularly during
hyperextension.
S Motion between two venebrae is small and
docs not occur independentl~y in vivo. Thus, the
functional motion of the spine is always a combined
action of sevel',,1 motion segments.
/6;- The instantaneous center of motion for the mo~
lion segments of the lumbar spine lIsually lies
within the lumbar disc.
The trunk muscles play an important role in
providing extrinsic stability to the spine; the ligaments and discs provide intrinsic stability.
8 Body position affecls the loads on the lumbar
spine. Any deviation from upright relaxed standing
increases the load. Forward flexion and simultaneous twisting of the trunk produce high stresses on
the lumbar spine.
9 EXlernally applied loads lhal are produced. 1'01'
example. by lihing or carrying objects Illay subject
the lumbar spine to very high loads. For the loads
Abt.:nhairn, L., Rossignol. ~'t., V~dal. J.P., l't a1. (2000). Thc r
of acti\'ily in the tht.:r:q1clllic management of back pa
Report of thl: Inll:rn:l1ion~1 Paris Task Force on Back Pa
5pinc. 25. IS.
Ad~lillS, M.A.• Dolan, P.. & 1·lutlon W.e. (1988). The lum
spin\.' in b~tChnlrd bl·nding. Spille. 13. 1019
Atbms. ~L\. & Hulton, W.C. (19831. The 1lll.:ch;lIlicill fu
tions of Ihe lumbar apophyst.\ljoints. Spil/e. 8, 327.
Alkn, C.E.L (1948). i\lusdc action potenlials used ill
sludy of dynamic analolll~', 13r J Phy.~ Jtcd, 11,66.
Anders()Jl, C.K., Chaffin, D.B., Hcrrin, G.D .. ct :11. (1985)
biolllcchanical model of the lumbosacral joint during l
ing ~\cli\'ities. J BiollH'dllllll·I.:S. 18, 57.
:\lHkl'sson. E.A .. Oddsson, L.I.E .. Grundstrom, 1-1., ct
(1996). E;\'lG acti\'ities or thl' qlladnHlIs lumborum ~
erector spinae musdes during flexion-relaxation and ot
motor lasks. CJinial! Bio1llcdUlllics, J J, 392.
Andersson, G.B.J. & l:l\·cnd"r. S.A. (1997). Evalu;ltion
Muscle Function. In J.W. Frymoycr (Ed.). Tilt' A(/1I11 Spi
Pril1c.:ip{es (lilt! Praclict: (2nd cd., pp. 341-3$0). New Yo
Li ppi ncot t-Ra\·l:n.
Andersson. G.B.J .. Ortengrcn. R.. 0:: Nachcmson. r\. (197
Quantitati\-c studies or b~lck loads in lifting. Spillt:. I. 1
Andersson. G.B.J .. Oncngren. R., & Nachcmson. r\. (197
lntr'ldiscal pressure. intr~\·... bdomin ... 1 prl.'SSUrl' ilnd m
det:tric bad lllusck :ll:tivity rdated to pOSl\lrc <lnd lo
ing. Clill OrtJlOp, 129, 156.
Andersson, G.B.J .. Onengl'cn, R., Nachclllson, A., et
(1974). Lumbardisc pressure and myoelectric b"H:k mus
~\clidty during sitting. 1. Stu,dics 011 an experimen
chair. SC(/I/d J Rclwhil ,Ht.:d, 6, 104.
,\sl11usscn, E. & Klauscn, K. (1961). Form and fUllction or
(.'rcc( human spinc. eJill OrlllOp, 25. 55.
Axlel". c.T. & McGill. S.M. (1997). Low back loads o\·cr'l \'
...' tv of abdomin.;\1 e.'\en.:ises: SeillThing for th..., s;lfcst
d~l1lin;l1 challeng(·. .\lcdiefl/I: & Sciem"":c ill Sports & E
cise. 29, 804
B:II·tclink, D.L. (1957). Thi.: role of i.lbouminal prC'ssure in
lieving the pressure on thl.' lumbar inu:n'crlcbral disc
BOlle }oilll Slfr~, 39B. 718.
B:ISIlWji.lll. J. V. (t958). Electromyography of iliopso'ls. :l
Rcf.:, 131. 127.
B~sl1laji~n. J.V. & DeLlll;;l, C.J. (1985). ;\1l1sdc:s Alive. Ba
mort: Williams &. Wilkins.
·f.
..;.'
,
BicrilH!·SorenselJ. F. (1984). Plwsical nh.'<lsurcmenls as ri~k
indlcalors for I()\\'-b~\ck lrol;bll.' over a on<:-yl.'ar period.
Spill!.'. 9, 106.
Callaghan. J.P.. P'IIla, .\..E .. & McGill, S.\\. (1999). Low back
thrcc--dimellsional joinl forces. killl.'malics. and kinetics
during walking. Clill;cal BioIJlC(://(lllics, I';', 203.
Cappozzo, A. (1984). Compressive loads in the lumb'lr vent:br.1! column during llorlllal le\'el walking. J Or/hop Res, l,
192.
Carlsoo. S. (1961). Thl.' st .. tic muscle load in differelll work
posilions: All dL'ctromyographic study. Ergollolllic.,·. .J.. 193.
Chaffin. D.B. (1969). r\ cornputcriz.cd biolllechanical modclDc\'dopmcl11 of and usc in stud~'ing gross body <lctions. )
BiOiI/Cc!wl/ics, 2, 429.
Chaffin, D.B.
Andersson. G.BJ. (1991). Occup(1tiol/al BioI/wcJ/(lIli<:s (2nd cd .. pp. 171-263". New York: 10hn Wile.\' 6.:
Sons. Inc.
Cheng. c.1\: .. Ch..:n. H.H .. Chen, C.S., et 31. (1998). Influcnces
of w~dking spce:d change' on the lumbosacral joint fon:c
distribulion. BiOl1lt'd Mara £11::', S, 155.
Choh:wicki. 1.. Juluru. K.. &: ~-IcGill, S.l\-l. (1999a). Intra-abdominal prt:ssun: ll1(.'chanisrn for stabilizing tll(.' lumbar
spine.) Bio/lwch, 32. 1-13.
Cholcwicki, 1., luluru, K .. Radt:bold. :\., I:t <II. (1999b). Lumb.ll· spinc stability (;'til be :lllgmcnted wilh an abdominal
bcli and/or increased inlra-abdominal presslIrt:. f.:'llropcall
'*
SJlill~ )ollrJ/al.
8, 388.
Choh.'wicki, J., P.mjnbi, ~L\L & f\.h;lChntr~·all. A. (199i). Sta·
bilizing function of trunk f1e.'l:or·('.'I:tcnsor llHlsch.. s around
<I neutral spine posture. Spine. 22(19), 1207.
Ciriello, \'_,\I....\: Snook, S.H. (1995). The dfl.-·Cl of b<lck belts
011 lumbar muscle fatiguc_ Spillt', 20, 1271.
COSSl.'ltc. J.W., Farfan, H.F.. Robertson, G.H., e[ at. (19il).
The instant~1I1l.'OUS <:enttr of rot~l[ion of Iht.' third Illlllb,lf
inll.'ITl.'nebral joint. J Biol/m.:It . .;.( 149). 19i 1.
Creswell, A.G. (1993). Rl.'sponses of intnl-abdorninal pressure
and ,Ibdomin~ll muscle activity during dynamic trunk loading in man.l::ur) Appi Physiol, 66,315.
Crt.·sswcll. A.G .. Blakt.'. P.L.. &. Thofstcnsson. A. (1994<1). The
drect of an ahdomin.d muscle Iraining program 011 ill[raabdominal pressure. Scant!) RC/Ulbil ;\ld. 26, 79.
Cr~ss\\"dl. kG .. Grundstr6m. 1-1 .• 6.: Thorslcllsson. :\. (1992).
Observations on intr<I-.llnlomillOlI preS$ur..: and p~lllcrns of
abdominal intra-muscular acti\'it~· in man . ..lcta Physiol
Scaml, 1';'4. 409.
Crcsswdl. A.G .• Oddsson. L.. & Thorstcnsson. A. (1994h). Thl.'
influcnce of sudden perturbations on ll'unk muscle: aCli\'ity and intrn.- .. bdominal prcssU!"t' while: sl<\nding. EXJll'l'
Brain Res, 98. 336.
Cresswell, A.G., & Thorstensson, A. (199-0. Chnngl.'s in intraabdominal pressure, Ll'llnk muscle acti\· ... tion and force
during isokinetic Iirting and lowering. Eur ] Appl Physiol.
68, 315.
Dalstr~l.
M. 6.: Huiskcs. R. (1995). Load lr.lnsfcr 'Kross the
pelvic bone. J Biollli:t:It. 28. i 15.
Davis. K.G., M ... ITas, W.S .. & Waters, T.R. (1998). Evaluation
of spinal IOildillg during lowering and lifting. CliJl Biolllt.'dumics, /3, 141.
de Looze, ~LP., Toussainl, 1·1.0.'1.. vall Dieen, J.H., et al. (1993).
Joint moments and llluscle activity in the lowel" e.'l:tremitics and lowcr ktck in lifting and lowering tasks. J 8iol/I.:c/',26, 1067.
Eie. N. (1966). Load capacity of the low b ..d. J Oslo Cit.\"
f1osp, /6. 73.
Eklwllll. J....\rhorelith. L·. F~lhlcr:Ill!I: ...\ .. l·t :11. (lYiY
\';lli(11I (If :,hd(llHin:lllllu:-clc:- during SOli\\,.' ph.\-sihl!ll,
til,.' c\l,.'l"ci .. l,.·!-. SnlJld) I?dwhil.l1t:d. II. 7.~.
101 U"hy. ,1.,.1, " K;n~, ..1.1. (1 'J~6j. Inlcn,'rlchr,d d;
f:l("l'! l,.·ollta("\ prl,.·ssufl· in a:\i~tll(lr!-i(1l1. In S.:\. l.anl
King (Ed:-.). /9SfJ :\d\'/I/H"".~ ill Uilh·lI:.;illt'I'l'ill.:: (pp.
:\l,.'\\' York: . \!l1l'l"icall SOl,.·iL·\\ or \In:h:\lliL':d Ellginl:(
r:ad~ln.ll.F. (19751. \ltlsl,.·III~II' Jlh: .... h;llli .. 1lI fir the lUJ11k
~tl\d I Ill,.' P(,:>ilioll of Pfl\\"l'!" and dficil·/lL·..... On/H
Xorth . \IlI. 6, 135.
Flinl . .\1..\1. (1965J. Ahdominal ll1u:>dL' in\'uln'l1ll'lll
thl' pl,.-rfnrm:lllt.·l,.· of \·.lriol1~ forms 01 :.it-up l,.·\l,.·n;
l,.·k('tr(Jm~·hgr'lphic :-llld~·. _'\1lI J Phys .\I,·d. ';'4. 114.
Fln\'d. \\'.F. <.\: Silv . . . r, P.fl.S. (193.:;). Thl' fUJll,.'linll (Ii lll
lors spinal,.' Inllsl·k" in cnl~lill 111()\clll . . . n\S ~llld post
man. J NI.niol. 129. 1$4.
Fri(·dd)old. G. (195~J. Dil,.· ..\kti\·iL'1 nOI·lIl.da Rih:h·n:
rnuskublur illl Ek,ktfom .... ogralllln ullll'r n'rschi
!·Iallllllgsik·dillgungl,.·n:
l,.·illc
Studil'
Skl·kllllltlskdnwchanik. I Of/hop. 90. I.
rukll\:tlJl:t, S .. :\,tbmur:l, T. nl·d~l. T.. l't ,d. (1995).
k,(t of Itlc(!1:ulic.t! slrl's;,,; 011 hvpntroph\' o!" 11k
li,\;;II11l'!lIUIll Il:l\'tllll. ) .')pilluIIJl'.-nn/. S. 126.
Gatlilk. J.O. (1<)67). h'n:>ilc propl'nil's 01 11K hUIll;11l
:lIl1lUIIlS fihrh:->II:-, .. lew O,.lIwJ' S,'lmd SII1'''!. 100, 1G;lrdll~'r·.\I(.tr~l,.·. .\I.G. & S!('Kl,.· .... LA. (19%l. Till,.' dfl'({~
dominal lllu:-dl' CI,·;ll·ti\:ttic.n on lumbar :-pilll' :S/Iill... II::!,'), ~9.
G(·rtl.lkill. S.D .. SI.·liglllt,IIl,.I. Hl.lltln, R....... l .d. (1%5
trndl' P~ltlL'l"llS :lnd sl'glllL'Il!,lI il1Sl~d)i1il.\ ill ,kgL'l
di:,\.· disl':l:-e. S"i!/,·. 10. 257.
Granat'l. K.P. .i: Sanford. :\.1-1. (19991. LUll1har-pd\'k
n~tli(ln i:> influl'IH..·l,.·d b~ lifting l~\sk par~\nh.'kr:>.
25(111.1413.
Grl,.'gCl":>l,.'Il, G.G . .i Luc~t~. D.l3. (1967). An in \'inl slud.\
lUlllbar a,xi:\1 rot;lliU!l (If thl' hlllll~\n Ilwrac(Jllllllb:\
.l 8011,' .luil!! Sur:: . .,19:\. 2-+7-232.
llaht.:r. T.IC. O·llrit:lI ..\1.. Dr.\·a, .1.\\'".. d 'II. (l\FHI. TIll
till' IUlllb.\I· f,ll,."l,.'! joinl::- ill spinal :-tahilil~·. IllcntiliGt
:illtTn:niH' palh~ of 1(1~\{lillg. Sl'ilh·. /9. 26-67.
Hannan. E...\., Fr~"klll:lll. P.:\ .. Clagclt. E.R ....... t ~tl. (19SS
,.hd(llllinal ,tlld in!r:l·[hor~lL'ii.' prt'-,,~uro during lift
jumping . .\led Sci S;10/"0 Erac 20. 195.
Hasq:;\\\:l. K.. T~\bhashi. IU_:. "Og~l. Y.. l,.'1 :1I. (199
Ch'Hlic:d propl,.,rti(·s or ostc'olwnic \,t:rtdn:d hodk's
torl,.·d hy ;Il,.·ou:-;tie l.'llIb:-ioll. /Jeme. /-1-. 7.'ii.
l-laugh!oJl, \'. .\1.. Schmidt. T.A .. Kl,.·l,.-k, K .. l,.'1 ;.1. (200())
bility of lumbar spin:d motion :-l'gllll'lltS I..'orrd:lll'd
of 1.:.lrs in Ilk' ;lIl11ulus fihro:-llS. J .Vel/rosur.::. 92. S
I-Iodgl,.·:-. P.\\'.. Crl,.·s:-\\dl. A.. l\: Thortl,.·l\:-~(JIl. :\.
Prl,.·p.\r:llor.\' [rullk lllUlioll ~ICc(lll1p:tnit:s r:lpid uppt
mO\·l,.'l1ll'l1l. 1:.\)1 Bmiu Ues. 12..+. 69.
Hodgl.'s, P.W. l\: G.l1HII..'\·j,l. S.c. (2000-,...\.t.·ti\·:lliO!l of
Ill,tn di:lphr~lgm during:l rt.·lk·tili\·l.' pOSIIlI';llt:lsk . .l
fiJI/I.!) . .)22. 165.
I-Iodgl,.·$ P. W. ~ Richardson C.A. (199iL Fl,.·cdforw:ml (
lion of tranSH'rSllS ahdomini:-; is nOI infhll,.·lln-d ll\'
l"l·...:lion of :lrm nW\·l,.·llH:nl. Ex!, U/"I/ill Ik~, 1/-1-. 36i.
I'[utton. W.e.
Adam::;, \-1.:\. (1982). Can tIll' lumklr IT\lshcd in ht:a\'~' lifting"? Spill!!. 7. 5S6.
'*
10s(·ph. J. 1.1960). "'flJI:': PO... ltll"t·: Ut!ClrolJl.'·fi~m/lhic S
Springfil,.·hl: Charll,.·s C Thollla~.
Jukn. D., .\!cGil\. S., l\.ropL P., ct ;11. ([ 99:n. OU:II11i[:t
tr:lllltl:-(lIbr llIyock·nril.' :H:ti\·it.\ of IlImb.lr porli
psoas and lhe abdominal wall during a wid\: \'ariel,v of
lasks. }fed S,:i SporlS f.XCfC. 30. 30 I.
A.I .. Pr:\sad. P.. ~ Ewing. C,1.. (1975). ~·h·dl<lnism or
spinal injur~" du(.' to l:<\udoCl.:phabd accderalion. Onlw/J
CIi" North :1m, 6, 19.
Kkin. J.A., Hickey, D.S., 0& Huskins. D.W. (I983). Radial
bul!:!in!: of lhe annulus fibrosus and the fUlll.:lion and fail·
llr~ -of ~he inlt:ITerlL'hr~t1 dbc. J Bi01llt'ch. 16. 211-217.
~'1.H .. S~:mllssi, R.E., Wilder, D.G .. el ,d. (1987). Inh:r·
nal displacelll~nl dislribution from in vilro loading of hu·
man lhoracic and lumb'll" spinal motion segments: Expl:r·
ill1l:ntal results and lhl'orelical predictions. Spill/.', 12,
1001-1007.
Krajc;Hski, S.R .. POI\'in, J.R .. & Chi;1I1g:, J. (1999). The in d\'O
r~sponse to the pl.:rlttrbations c:lllsing rapid f1cxion: Ef·
f~cts of pre-load and step inpul magnitude. Cii,l BiO/lll/dUlU;CS. 14, 54.
Kuhlk. R.E. Schuhz. A.B .. BdYlschko. 1., el <II. (1975). Bin·
ll1echank;tl charactcristics of \'crICbral motion segments
and intern:rtcbral discs. Or/hop Cfill North AI/I. 6, 121.
L1nd~r. J.E .. B~ltes, B.T.. & Devita. P.(1986}. Biolllt.'chanics of
th\.' squat (.'xen:isc using. a modifil.'d Ci.'Ill(.'r of mass bar.
.\lcd Sci Sports E.un:, IS. 469.
Ll\'cnder. $.:\., ;\lirka. G.:\ .. Schol.'I1ITI.lrklin. R.\\'., el al.
(1989). Thi.' l:rrCC1S of pre\'iew :lnd t<l.sk sym!lll:lry on trunk
mllscli.' response to sudden loading. lIwlIlm Fat-tors. J I,
101.
[.:\\'coder, S.A .. Tsuang. 1".1-1 .. AIH.li.'rsson, G.B .. ~t a1. (1992).
Trunk nlusch.' co·contraction whitt: resisting applied mo·
ments in a Iwisted poslure. Ergollolllics. 36, 1145.
Lt.'c. Y.H. &. Ch('ll. C.Y. (1999). Lumbar vcrtebr.d angles and
b.lck muscle loading with bells. Illd Health. 37, 390.
Lucas. D.B. &. Brcsier.~B. (1961). SIal,ilil.'· of ,ht: Ligamt.'I1fOUS
SpiltE.'. Biomechanics Laborator~', University of Cnliforni:\,
San Francisco and Berkeley. Technical Repon 40. San
Francisco: Th~ Ltbor.llor\,.
Lumsden. R.M. & Morris. J.~'1. (1968), :\n in dvo study of axinl rotation and immobilization at the lumbos:\cral joint.}
Heme Joill/ SlIrg. 50A. 1591.
Manas, \V.S. & Granata. K.P. (1995). A biomechanical assess·
ment and model of axial t\...· isling in the thoracolumbar
spine. Spil1c? 20, 1440.
:\larras. \V.S. & Mirb. G.A. (1992). :\ comprehcllsivl.· c\'alll:\lion of trunk respon$c 10 asymmL'tric trunk lllOlion. Spill/!.
17. 318 .
.\larras, \V.S .. Rang<lrajuili. S.L.. & Lan'nder. S.:\. (1987).
Trunk loading and ('xpectation. Erf,:ollolllics. 30, 351.
McGill, S.~vl. & Norman, R.\V. (1987). Reasscssm~nt of the
role of inlra·abdominal prcsstlrt.' in spinal compression.
Ergollomics. 30, 1565.
McGill. S.M .. NOrllmn. R.W., &. Sharratt. M.T. (1990). The d·
fret of an 'Ibdominal b~1t on trunk muscle acti\'ity and
intra·abdominal prl.'sslln..' Juring squal lifts. Er~ol/olllic.s.
33, In.
McGill, S.~1.. Yingling, V.R .. &. Pl..'ach, J.P. (1999). Tlm..'I..'dimcnsion:d kin~rnatics and trunk muscle my()d~ctric ac·
ti\'it~' in the ddcrly spinc-<l dawb,lse compared to young
people. CJiIl Biolllcclu/llics. 1'+. 389.
,\'111es, M. & Sull ivan. W. E. ( 1961 ). L:lti:nd bi.'llding .11 Ihe hlln·
bar and lumbosacral joints. .·\11111 U~C, /39. 38i.
~liller. l.A.A., Hadersp\.'ck. K.:\., 6.: Schuhz, A.B. (1983). Pustl'rior (.'lelllenl loads in lumbar motion segmcnts. Spille. S,
331.
~Ioll.
.I.M.H. &. Wright. V. (1971). Normal range of spin;.
bility. An objcclh'e clinic:l! study. AIlII UhClIIll Dis, 30
Morri!'O. J.!"..l.. Bl:lllli:r. G., ~ Lucas. n.B. (1962). An
tf(Jl1lyographic study of Ihe intrinsic muscles of the
in man. J :111t1l l.mul. 96. 509-520.
~-lorris. J.~L, Lucas. D.B .. &. Bn:sier, B. (1961). Role o
trunk in st.:tbility of the spin<.'. } BUIIi: JOil1f Slll~~ . .:31\
:\-losekilck. L. (1993). Vertebral struclllr\.' :lIld strength in
and in vitro. Cah:i{Tissllc 1111.53.5121.
N;'lChl'llIson. A. (1960). Lumbar inlrndisc'll pn:SSllrl.' ..-k/
[hop Sculld SlIp,.,l.
n.
1-140.
Nachclllson, A. (1963). Thl' influence of spinal rno\'erni..'
till.' lumbar intrndiscal pri.'SSUl'l.' and on till' tensile str
in IIIi.' ;tnlltllus fibrosus . ..lela Off/WI' SClIIul. 33, 183.
Nachcmson, :\. (1966). EIeCiromyographic studies Oil th
td}ral portion of the PSO:IS llluscl~. With special rcr
to its stabilizing function of the lumbar spine. lkw O
Scalld, 37, Iii.
Nach<.:mson, A. (1975). Towards a beller understandi
back pain: A review or the mechanics of the lumbar
RlJellllwlO! RelUlhi!. I.J, 129.
Nach(.·m;.;on, A. & Elfstrom. G. (1970). IlIlravital [)\'I
Pre:ssure: .\lC(ISllrCJ1h.'lllS ill Llflllh(lr Dis,;:,: ,.\ Sllui.\' r.;{
IlIOIl .\IOl·i'J11CIlfS. :\Ialli:III'crs allll Exercises. Stock
Almquisl & \\"iksc-lt.
~achelll;.;on, r\.L. .& E\':lns. J.H. (1968). Some mech
propcnies of the third human lumbar intalaminar
ment (IigameJltulll flavum). J BiomcclulIIh's. I. 211.
Nachclllson, A. & Morris, J.\l. (1964). In \'in) mCaSUri.
of inlradiscal pressure. Discolnl'lry. a mcthod iol' th
termination of presslll"i.' in the lower lumbar discs. J
1ui1l/ SI/rg, -16:\, 1017.
!\alional Institute for Occupational Safety .tnd Health (1
WOl'kplclCe Use
S(lck Be/IS. DHHS (NIOSH), Nu
94-122.
Ndson, J.M .. \blmsky, R.P.. 6.: St(.'n:nsoll, J.,\'I. (1
RcI'lti,·c IUlllbm' ;tnd pl.'!\·ic motion during londcd
fk'\ion/extcllsion. Spinc, 20. 199.
Nemeth. G. (1984). On hip and lumbar biomechanics. A
of joint load and muscular ;lCtivity. SC(/l/(! J Uelwhi
or
Stlppl, 10.
Nonon, P.L. & Brown. T. (1957). The immobilizing c-ffi
of h'ICk braces. Their eHecl on th~ posture and mot
thl..' 11l1lloosacr;.\1 spine. 1 BOIIt' }oill1 S/lrg. 39:\. Ill.
Dncngrcn, R. And~rsson. G.B.l., & 1'\achcmson. A.L. (1
SllIdies of rd:lliollships bel'\'cen lumb,lr disc pre
myoelectric b<'H:k muscle acti\'it~,. and intra-abdominn
trag;:\slric) pr~SSllre. Spille, 6. 98.
Osval<k'r. A.L.. Ncumann, P., Lo\"sund. P.. et a1. (1990)
maII..' strength of the lumb.\r spine ill fle.xion-.1I1 in
study.} Bio1l1t.:clull/ics, 23, 453.
Panjabi. ,\13\-1., Gael, V.I-\. .. ..\: Tabtn, 1-\.. (1982). Ph~'si
strains in the lumbar spinal ligamC'1l1s. An in vilro b
chanical slUdy. Spi1/i', 7. 192.
Parnianpour. M., Nordin, \1.. K<thnnol"iti'.. N., l.'t al. (
The tri<.lxia! clHlpling or torque gcneration of trunk
cles during isometric e.'\cnions and the effect of f:tl
isoincrtial mo\,emenlS on till.' motor output and 1ll0\
p~IIt('rns. SpiJlt', 13(9).
982-991.
Partridge. M.J. 6.: Walters, C.E. (1959). ~articip.llion
abdominal IlHlsdes in \'arious mOVt'lllc-lltS of Ih ..~ tru
IlWIl. An L'!cl::trol11yographic study. Phys Tll!'r Rei', 39
Paul.\', J.E. (1966). :\n dcclromyographic '\llaly~is of cl:nain
nlO\'('Il1(:nls <lilt! ex('rcises. I. SOllle c!('cp muscles of Lhl'
f
back. All(/( Ret:. /55. 223.
Perkins, ~I.S. & Bloswid. D.S. (1995). Th,· usc of h'ICk belts
10 increase inlr~wbdominal pressure as a means of pn:venting lo\\' back injuries: A survey of lhe lih:r.lIun:. 1m J
OCCIlP EUl'imll rlt.:lllih, I, 326.
Pope, M.H., Andersson, G.B.J., Brolll.lIl, 1-1., L't .d. (1986).
Electromyogr:q1hit studies of the ]lImb'H' trunk muscula.
tun: during thl: dcvelopmcnt of a.xi.1I torques. J (Julio!)
Res, cJ, 288-297.
Ranu, B.S. (1990). Measuremctlt of pr(.'SSllrcs in the nucleus
and within the annulus of the hurn'lll spinal disc due 10 l.:X'
trcme loadilll.!. Pmc Illst Jleclt
Rcichm;t1Hl,
S.~ (1971).
Ef/J,!,
(llJ. 204, 14l.
Motion or thc
lumbar anicular
processes in fh:xiull-cxtcnsion and lateral flexion of the
spine. ~1c[(t .\!orp!lo/ Nt'crl SCfllld 8, 261.
Reichmann, S., Bl.'rglllnd, E.,..\: Lundgr(,.·n, K. CJ972). Das B(,.·.
wegungszcntrum in der LendenwirbclsAule bei Flc.xio[l
lind Extension. Z A/llIl EIII\\·ickltlll;;';;;':I!.\·(.·h. /38,283.
Reid, J.G. &. Costigan, P.A. (1987). Trunk llluscle balance and
muscular force. Spill/!, /2, 783.
Reyna, J.R. Jr., Leggdl, S.I-I., Kenney, ~., "I .d. (1995). The ef.
feCi of Ilimbar bells on isolated IlImb,lr musl'!e. Slrength
and dynamic (;"padt)'. Spilll.', 20, 68.
Rolandcr, S.D. (1966). ~'lotion of lhe lumb'lf spine wilh special reference to the stabilizing effect of posterior fusion.
An experiment"l study 011 autopsy specimens . .'\elll Ortho}l
8colld, .)11]1]11. 90, 1-144.
Sato, K., Kikuchi, S" & '{onez;Jwi1, 1'. (1999). In vivo intradis.
cal pressure llle'lsuremCllt in healthy individuals and in
p<iticnts with ongoing back probll.'ll1S. Spille, 24, 2468.
Schuhz, A., ~l ~\1. (1982). Loads on lhe lumbar spine. Valida.
lion of a biolllcchanicaJ analysis by mC<lsurenH.'IllS of in.
Ir~ldisc.d prc:sslIn:s <1I1d myol.'!l.·ctric sign ... k J BOIlt' Joil/!
Surg. 64.·\,713.
Shir4lzi-t\dl. ,.\. (1994). Biomt'ch4lnics of lhe lumbar spine in
sagittalJlatcral moments. Spi/h:, /9, 2407.
SIlI(ImOIlO\\" \\., ZIHIlI. B.-II.. Bar:llia. R.\'., 1.'1 itl. (19
llh.'l'h:lnic. tlf incrl'a:-I.:d expUSlIl"(,.· to hlilihar inju
h.l.· l'~"l'lic loadillg: Pan I. lo;.;s of rdl,:.\i\·(,.· 1Il1l:,(,:
liz;llioll. Spillt'. 2·H23). 2-l26-243-l.
Sldkn. T., B;II·alllki. II.C .. Ruhin, R.. d :11. (19t.1~J_
inlradisl'al pr(,.'~:-.ur(,.· Ilh:a::-urcd in I hI.' :I11t(,.·riol" itllt
bler:!l anllul;,r rcgi(l/IS during :i='.\"Inml.'lricll I(I;l
Biol/f(·clu/!/ic.\. 13, 495.
Sllln:ssoll, B" Sl'h·ik. Go, & L"dCIl, A. (1'J89). i\!on
thl.' s:lcroili:ll' joints. ,\ rt)I.'rllgt·ll sLerl'oplHllOg
:lIwlysis. Spi!lc, 1-1, 162.
Thomas, ./.5., Lavelltk'r, S.:\ .. Corcos, D.M., 1.'1 :t1. (1
fl'ct 01' lifting bl.'l!:- Ol1 ll'unk Illusck :1l"lil':Lli()[l
sudlk'111.\" appli<:d In.td. /l1t'WIII Faclo!"s. ·iI. 6iO.
luong, N.H., {),lll:-l,.·r(,.·:HI. J.. \latlr~tis, G.. 1.'1 ~d. (199S
dillll..'llsional I.·\·alu.ltion of lumhar urthosis d
:-pinal bL'h;l\"i\11". J Uclwbil Il.·... Del'. 35. 3-1.
lrb'lIl, J.P,G. & :\kMllllin, J.r. (1985). Swdlillg prl
thl' inleJ"l.·crtd.u·al di!'L'; Infllh:llc(,.' 01 pmIL'oglyc;1I1
I:q;l'n COlllt:'IlI:-. Bifl,.h.,o!o.t:..\·, 22. 1-45.
\',111 Dici:n. l,tI., llUCt"l.I.'lllan:-, M.L\!.. &. TUllssai
(1999J. SH,op or squal: ,.\ rl.'\·it'\\· of bilJlll(,.·<:hallil:a
on lifling 1I.:(hniqu(,.'. CliJl Bifl/lft'c}wuic:•. /4. 6$5.
\bn~..\1. & DUlllHS. G.:\. (199S). \kdl;lIlh,:al hdl;l\'i
fl.~lllak s:l(roilial.· joinl [llld inflllL'nCt' of t!l(,., :l1l11
poskrinr s:II.::roilia<: Iigamcnts under sagittal In
Hiw!lccltullio. /3, 193.
\\'hi!l.', :\..-\. (1969). An:d.\"sis of thL' Illl'I'!l,lnit:- 01'
spint:' ill Illan. An expl.·rillll.'llt,t1 stud.\' (II' :lllIOPS
mens . .·le/ll ()rt/wp ,"jcl/ud, SIIp/i! il7, 1-105.
Whilt:, A.A. l\:. P:lnjahi. :\I.N. (19701. Cliniud HiulIlt'cI
Iht'Spilit'. I'hil:ldl.'lphia: J.B. Lippincott.
""ilkl.', "1.1., Ned. Po, C:lillli. :\1.. <'I al. (19991. ;\I.'W
ml.·asur(,.·IIK·nts of pr(,.'ssurl.'S in Ihl.' inl<.'l'\·(,,·rlChra
dail~' life. Spill.·, 24. 755.
Wilder. D.G., POpl.· ..\1.1-1 .. &. Fr.\"Ino~·L'r. J. W. (1980). T
tional lopogr,tphy of the s;Jcroiliac joint. Sp;/It'. 5,
Biomechanics of th
Cervical Spin
Ronald Mosko
Introduction
Component Anatomy and Biomechanics
Anatomy
Osseous Structures
Intervertebral Discs
Mechanical Pmperties
Vertebrae
Intervertebral Discs
Ligaments
Muscle
Neural Elements
Kinematics
Range of Motion
Surface JOint Motion
Coupled Motion of the Cervical Spine
Atlantoaxial Segment
Subaxial Spine
Abnormal Kinematics
Spinal Stability
Occipitoatlantoaxial Complex
$ubaxial Cervical Spine
Applied Biomechanics
Decompression
Arthrodesis
Cervical Spine Fixation
Biomechanics of Cervical Trauma
Airbag Injuries
Whiplash Syndrome
I·
•
:1(
il1Lh"'d,~.,%>. mLdi.
Summary
References
Introduction
Knowledge of spinal biomechanics advanced exponentially during the second half of the twentieth
centlll)'. A two-column model of the spine was describcd by Sir Frank HoldswOIth (1963) and, latel; a
lhree-column model by Denis (1983), further refining the principle or spinal stability. The computer
age has produced powerful methods For modern
hiomcchanical modeling, the promise being the
ability to assess the slnbility of n construct prior to
implantation. Today. the application of cendcal biomechanical knowledge spans many inclustl'ies and
supports improved medical diagnoses and treaLment that is more effective. Future technological
and electronic advances will continue to build on
basic biomechanical principles, many of \\:hich will
be outlined in this chnpter.
arc five t.'·pical cervical vertebrae, C3-C7, which arc
sirnilar in structure and function.
The spine has four curves when viewed in the
sagiual plane. The celvical and lumbar regions arc
convex anteriorly (lordotic), while the thoracic and
sacra) regions are convex posteriorly (k)iphotic).
The lordotic curves develop after birth as the infant's spine straightens out, which facilitates development of a bipedal posture. Although there is a
Cl (atlas)
C2 (axis)
Facet
joint
Component Anatomy and
Biomechanics
C7
ANATOMY
The exquisite design of the cervical spine uniquely
contributes to the structure of the human body and
profoundly enhances its function. The cervical
spine supports the skull and acts as a shock absorber for the brain. It also facilitates the transfer
of weights and bending moments of the hend. It
protects the brainstem, spinal cord, and various
neurovascular SU'UClures as they transit the neck
nnd when they enter and exit the skull. The vertebral column also provides a multitude of muscle
and ligamentous attachments for complex movement and stability. The neuromuscular control afforded by the muscle ntlnChmenls combined with
the numerous articulations of the cervical spine allows for a wide range of physiological motion that
maximizes the range of motion of the head and
neck and serves to integrate the head with the rest
of the body and the environment.
The spine consists of 33 vertebrae divided into
five regions: cervical (7) (Fig. II-I), thoracic (12),
lumbar (5), sncml (5 fused segments), nnd coccvgeal
(approximately 4). The two most cranial vertebrae,
Cl (atlas) and C2 (axis). [Ire atypical. with a unique
structural role in the articulation between the head
and the cervical spine. The atlanta-occipital joint,
bel ween CI and the oecipital bone of the skull, is
also a functional part of [he cervical spine. There
A
Uncovertebral
joint
Intervertebral
disc
Tracheal
air shadOw
C7
B
A, lateral roentgenogram of the cervical spine. Note the
lordosis. The facet joints are aligned obliquely only to the
frontal or <oronal plane; hence, their excellent visualization in the lateral view. a, Anteroposterior view of the <er·
vical spine.
287
•> QJ- ._
harmonious progression of these curves from one to
another, which may help distribute stresses and
strains, injuries occur more commonly at the jllnc~
tional areas because of differences in the relative
stiffness of each anatomical segment of the spine.
The physical structure of the anatomical elements
modulates from the cervical to the sacral region in
relation to the segmental function.
The lordosis in the cervical spine, like that in the
lumbar spine, is maintained predoI1linantl:v b.v
slightl)' wedge-shaped intervertebral discs that arc
larger anteriorly' than posteriorly. In contradistinction, thoracic kyphosis is maintained largely.' by the
vertebral bodies themselves; because the posterior
portion of the thoracic vertebral body is larger than
the anterior portion, there is a relative kyphosis of
the thoracic spine.
The conceptual biomechanical building block of
the spinal colunln is the Functional spinal unit or
motion segment. It consists of two adjacent vertebrae and the intervening intervertebral discs and lig~
aments between the vertebrae. These ligaments arc
the anterior and posterior longitudinal ligaments;
the intertransverse, interspinous, and supraspinous
ligaments; and the facet capsular ligaments. As a re~
suit of the different functional demands of the various parts of the spinal column, segmental variation
is expressed by! changes in the size and shape of the
vertebrae, the anatomy of the discoligamentous
structures, and the alignment and structure of the
facet joints.
Biological structures behave differently than do
common engineering materials. Collagenous tissues
exhibit both viscoelastic and anisotropic behavior.
Viscoelastic properties are rate~dependent (timedependent) behaviors under loading that are seen in
both bone and soft tissues; mechanical strength increases \vith increased rates of loading. Anisotropy
is the alteration in mechanical properties that is
seen \vhen bone is loaded along different axes.
Anisotropic behavior occurs as a result of the dissimilar longitudinal and transverse microstructure.
sagittal range of motion of the cervical spine.
Cl-C2 articulation is the joint primaril.\-' respon
for rotation in the cel'Vical spine.
The atlas, or C 1, is a bon.\-' ring consisting o
anterior ;:\lld posterior arch that is attached to
two !;:\terallnasses of the atlas (Fig. 11~2). Th
pcrior surfaces or the latent! rnasses, which
crani<:dly <:md inward, form an <:\rticulation with
caudally and clutward-facing occipital cond.\-'l
the skull (Fig. 11-3). E.xtension ()f the occipit
vical joint is limited by' the bon.\-' <:uwt()!ll.v; fle
is lirnited prirnaril.\-' by ligarneluous structures
tectorial mernbrane, and the longitudinal fibe
the crucifonll ligarnent as well as by the post
ligaments (Figs. 11 ~4 and 11 ~5). The anterio
berek' on the <:llTh of C I serves <:lS <:In attachm
for the longus colli muscle, a flexor of the n
The posterior arch of the atlas is a modified lar
that is grooved on its superior surface for the
sage or the vertebral arteries as the.v enter
the fClrarnen magnum after piercing tile post
ntlanto~(lccipital nlelllbl"ane.
Simi!;:\r to the occipitocel'Vical junction, the
11() intervertdJral disc betwecn Cl and C2. Stab
at this levcl is thus predicated on intact ossco
l1lentclus structurcs. The articulation between
and C2 is primaril.\-' specialized for rotation.
bod.\-' of C2 projects superiorly to fOlTn the odon
process, or dens (Fig. 11 ~6). The projection o
bod.\-' or C2 and the dens has a characteristic ob
appearance on lateral cervical radiographs an
Superior articular
facet
Orig
tran
liga
Groove
vertebral artery
Osseous Structures
The occipul-CI-C2 complex comprises the upper
cervical spine and is responsible for approxilnately
40% of cenrical flexion and 60% of cervical rotation.
The occipital condyles articulate with the slightly
concave lateral masses of the atlas. The primary
Illotion permitted by this articulation is flexion and
extension, accounting for a large portion of the
Facet for dens.
on anterior arch
1cm
I
on posterior arch
DmI'----
1
Bony architecture of the atlas. Bar
" "__" ~
1 C111.
Occipital
condyle
,r--__
Occiput·C1
joint
Dens
Transverse
process
Cl-C2
facel joint
Lateral mass
of Cl
Lateral mass
ofC2
.
The open-mouth radiograph demonstrates the occiput-(l and the atlantoaxial articulations. Note the symmetric spacing between the lateral masses of (1 and the dens. Asymmetry or widening of these spaces may occur after rotatory disturbances or fractures of
the (1 ring .
--------------------------
Anterior
I
Apical
ligament
01 dens
Posterior
Superficial layer of
tectorial membrane
Tectoria! membrane
! Ant:~~~
I of alias
I
",,","u
~-~
Dens of axis -t;~ffB~ll
Transverse
ligament ~~~~y..
.
~~
)'.,
-.:,::(-'\
~..
Anterior
I atlantoaxial
I
I
@0
ligament
Flexion
interv~~:bral
fibrocartilage
Ilongitudinal
Anlerior
1
ligament
Body of C3
Posterior
longitudinal
ligament
Median sagittal section through the occipital bone and the
first three cervical vertebrae showing the articulations and
surrounding ligaments.
Extension
Tracings of lateral flexion and extension radiographs
ing the occiput, el, (2. and (3. The substantial relati
motion between the occiput, (1. and (2 can be seen
Large arrows indicate the direction of motion_ Small
rows indicate that approximation of the posterior ele
ments limit occipitocervical extension. In contradistin
maximum flexion is controlled by tautness of the Iiga
ments. Reprinred with permission {rom MoskoviclJ, R. & J
D.A (1999). Upper (eNica/spine instrlJmellt<11ion, Spine:
of tIle Art Reviews. 13(2), 233-253.
disc, The
Dens
(odontoid
process)
\'crtebral hod,\" is o\'ul-slJaped
transverse pnlcL'ssCS
Spinal canal
Superior
Spinous
facet
process
Foramen
lransversarium
cCryiC~IJ
is widl.:l" mediolaLerally Lhan <:Illtcropostcriorl:",
Inferior facet
lcm
or the cervical spin\,.' arl' uni
in thaI Lhe,Y all contain a lr~Hlsn:rse foramen for
passage of thc vertebral a ncr:', Tl\\"~ transv
processes or the suba:xial c(~r\'ical \'('ncbr~\e h
two projections, lhe alllerior and p(Jstcl'ior tu
cit's. which serve as allilChl11\,.'nt points for ante
and poslcriOl- lTluscles, respectively, The large a
rior tubercle of C6, rt'fcrrcd to as the carotid tu
cle, can be an important surgicallandmal'k, The
peric>!' surract;.' 01' the lrans\,\,.'l'se proc~ss provid
groove for the exiling nerve root.
Each pedicle connecLs the vl".·ncbral bod~! to a
eral mass, that portion of bone containing lh~ s
rior and inferior facets, The facct joints n.::gulatc
or
mO\'l".'ll1ents
the spine and pla~- a c.:ritical rol
spinal slabilit~" Those 01' the cervical spine are orie
The axis vertebra, or (2. The superior facet articulation
permits multiplanar motion while the inferior facet is
aligned to articulate with a more typical cervical facet,
which is more constrained. There is a smooth surface on
the front of the dens for articulation with the anterior
ring of (1. Bar>:: 1 em.
at approxirnatd~' 45'~ to the coronal pbne and arc
cated in the sagittal plane (Figs, 11-8 and 11-9),
orientation allows grc(l(('r anlounts of flexion t
does latl".'ral bending or rotation in the cl.;.'lyical sp
The facet joims resist most of lhe shear forces and
often a helpful anatomical landmark. The dens ar-
proximalL'!:' 16l.:!r; of lhe compressi\"c forces acting
the spine (Adams & I-lutton, 1980). The laminae
arise from the lalcral masses, Tht' IaLeral rnHSS\,.'S h
important surgical implications in tht: subaxial ce
cal spine because they conwin a relativel:' l
ticulates with and is restrained within a socket
amount 01" bone and are easil~' accessible for the pl
•
formed by the transverse ligament of CI and the anterior arch of C1. The transverse ligaments run from
the anterior arch of C I. behind the dens, and prevent anterior translation of Clan C2. The other ligaments at the CI-C2 articulation are the alar, apical,
and accessory alar ligaments. The alar ligaments,
which are symmetrically placed on both sides of the
dens, attach the dens to the occiput to prevent excessive rotation. The left alar ligament prevents
right rotation and vice versa, To some extent, the
alar ligaments also act to limit motion during side
bending (Dvorak & Panjabi, 1987). The apical ligament also connects the dens to the occiput.
Unlike the 1\\'0 most cranial vertebrae, the
anatomy of the third through the sixth cervical vertebrae is similar (Fig. I 1-7). These fOUl" cervical vertebrae consist of a body, two pedicles, two lateral
masses, two laminae. and a spinous process, The
seventh cervical vertebra is slightly different in that
it has a transitional form, It is called the vertebra
prominens and has a larger spinolls process that is
not bifid like those of C3-C6.
The anterior components of a subaxial cervical
motion segment are the vertebral bodies and the
Inferio
facet
Groove
for spinal
nerve
Super
facet
Pos
tube
Anteri
luberc
Uncinate
process
lcm
Superior view of a typical cervical vertebra, representa
of (3-(6 ((7. the vertebra prominens. diHers slightly in
that it has a prominent nonbifid spinous process),
C4
CS
C6
C7
Orientation of the facets of a typical cervical vertebra in
three planes. The facets are oriented at a 45" angle to the
transverse plane and the frontal plane, and are at right
angles to the sagittal plane. Y indicates the craniocaudal
axis. z the anteroposterior axis. and x the mediolateral
axis. Adapted '.t'lieh permission from White, A.A. 111 & Panjabi,
Lateral photograph of a fourth and seventh cervical ver
bra. The facet joint alignment is fairly close to 45° from
the transverse plane. Note also the difference in size of
the spinous processes, which are a reflection of the size
and importance of the muscle attachments_
M.M. (1990). Cfinical Biomechanics oi the Spine. Pili/adelphia:
I
1. B. Lippiflcort
e
menl of screws. as opposed lO the pedicles. which arc
difficult to cannulate sarelv in the neck.
The superior surfaces of the cervical vertebrae are
saddle-shaped because of the uncinate processes,
bony protuberances that aJ-ise from the lateral margins or the superior end plates (Fig. I 1-10). The uncovertebral joints Uoints of Luschka) develop during
spinal maturation and pla~,/ an important biomechanical role with respect to kinetics and stability.
•
Intervertebral Discs
The intervertebral discs are highly specialized st11.1Ctures that contribute up to one-third of the height of
the vertebral column and fmm specialized joints between the cartilaginous end plates of the adjacent
vertebral bodies. Activities such as running and
jumping apply short-duration, high-amplitude loads
to the intervertebral discs, whereas normal physical
activity and upright stance resuh in the application
of long-duration, low-magnitude loads to the disc.
Discs are able to withstand greater than normal loads
when compressive forces arc rapidly applied based
on the biomechanical principles of viscoelasticity.
This propeny protects the disc /i'om catastrophic
failure until extremely high loads are applied.
1cm
Anterior view of a sixth cervical vertebra. The shorr arro
indicate the uncinate processes and the Ic;mg arrow ind
cates the pathway of the sixth cervical nerve root. The
facet joints are located posteriorly. Bar = 1 em.
The nucleus puiposlis is centrally locmed within
the disc and consists of almost 90% water in young
individuals. The water content is highest at birth
and decreases to approximately 700/0 as the disc degenerates with age. The rest of the nucleus pulposus
consists of protcoglycan and collagen, which is exclusively type II collagen. Type II collagen fibrils are
thought to be able to absorb compressive forces better than type I collagen fibrils.
Prolcoglycans consist of a protein core auached
to polysaccharide (glycosaminoglycan) chains. The
pol.vsaccharicles are either keratin sulfate or chondroitin sulfate. The core protein, with its attached
polysaccharides, is aggregated to hyaluronic acid
through a link protein. The proteoglycans in th(: intervertcbral discs are similar to those in articular
cartilage, except that the protcoglycans present in
[he intcl\'ertcbral discs have shorter polysaccharide
chains as well as shorter core proteins. The nucleus
pulposlls contains more protcoglycan than does the
annulus fibroslis. \¥ith increasing age and disc degeneration. the lotal protcoglycan content decreases.
The annulus fibrosus is the OtHer portion of the
disc. Its water content is slighl1y less than that of the
nucleus, being only approximately 78£10 \vater in
younger individuals. \-\lith age, the water content
falls to approsimalely 70%, like that of the nucleus
pulposus in older persons. The annulus consists of
collagen that is arranged in approximately 90 concentric lamellar bands. The collagen fibers in these
shccts nll1 at approximately 30 0 to the disc or 1200
to each other in the adjacent bands. This unique orientation confers strength to the annulus while permilling some nesibility (Fig. I I-II). The composition of the collagen in the annulus is approximately
60% type II collagen and approsimately 40% type I
collagen. As the disc ages, the collagen undergoes
irreducible cross-linking and the relative amount of
type 1 collagen increases, replacing type [[ collagen
in the disc.
MECHANICAL PROPERTIES
Vertebrae
The mechanical properties of the bone and soft tissues diffe!: Strength and stiffness and the relation of
stress to strain are the key IT'lCchanical properties.
Stress-strain curves are used to determine the relative loading behavior of bone. Stress is the load per
unit area of a perpendicularly applied load. Strain is
[he change in length per unit of original length, usually expressed as a percenlage.
:lj
,
I
.ZiZY>_
.~"
).
-~-"'>--
-
~."
.
~~~~~~~~?-I)
,./ pulposus
Nucleus
(
A
I
Disc -
An
fibe
'
~~~
I
Schematic drawings of an intervertebral disc showing
criss-cross arrangement of its fibers. A, Concentric lay
of the annulus fibrosus are depicted as cut away to s
the alternating orientation of the collagen fibers. B,
layers of the annular fibers are oriented at a 30'" ang
the vertebral body and at 120 angles to each other.
Acf,;pred '.-Vlil! permission l,om W;uw, AA. 111 c~ Pan,iabl. M
0
(l990} Clinical
B!Drn~'dl,Hll(<'
of the Sp:rl(' Pili/dcle/ph/a: I
Lippincor[
!
•
Cortical bone is stilll.'!" than cancellolls bone
can withstand greater stresses before failure. \V
the strain in "h'o excl.;cds 2t:'(; of the originallen
conic::d bone l"1"<.tclLtrcs, but c<.tllcdlous bone
withSland somewhat greater strains before fra
ing. The gr('.. ltl.~r abilit~· to withstand strain i
cause of" the stl"uctUrt: of cancellous bon ... : its p
it,v varies from 30 t() 9CY>" compared with con
bone with valuL's or 5 to 30%; (C~\rtcr & H
1977). Vertebral corn pression strength incre
from lhe upper cervical lo th(' lower lumbar k\
Bone strength decreases with age as a resu
[he de\'c1oplllcnt of osteoporosis. The mineral
tenl (If' vertebr~\e decreases with increasing age
relatively con::'1ant rale (Hansson & RODS,
Hansson et 411., 1980).;\ 25 cj{; «(..crease in osseou
suc results in a 1110re lhan a 500'0 tkcrcasc in
strength of the \"Crlchrac (Bell. 1967). BeC'lllSC
cortical shell of a venchra is responsible for on1
proximatcl.v 1(Y;',; 01" its strength during com
'sion, good quality c~lIlcellotis bone is criticall~
!Jorlant (McBroom e:t aI., 1985).
\Vht,:.~n bone is loatk.'d in "h'o, contraction o
musclt:s attached [0 the bone can alt0r [he s
distribution in the bone. Bending moments are applied to the vertebral bodies during motions. During
lie.""", tensile stresses are applied to the posterior
and compression to the anterior cortex of the
body. To perform lifting tasks, the back
are required to develop considerable forces
IS"hltltz et aI., 1982). Stresses in a typical cClvical
change from tensile to cornpressive in a region approximately 0.5 to I cn1 anterior to the pos~
longitudinal ligament (Pintar et al., 1995). As
is weaker and fails earlier in tension than in
compression, posterior paraspinal muscle contraccan decrease the tensile stress on bone by! producing a compressive stress that reduces or neutralizes the posterior cortical tensile stresses, allowing
the vertebrae to sustain higher loads than would
othenvise be possible. However, bone will often fail
prior to damage occurring to the intervertebral disc
under compressive loading. Finite element modeling of the cervical spine indicates that the increase
in end plate stresses may be the initiating factor for
failure of this component under compressive loads
('loganandan et al., 1996a).
Intervertebral Discs
Intenlertebral discs exhibit viscoelastic properties
(creep and relaxation) and hysteresis (Kazarian,
1975). Creep occurs more slowly in healthy discs
than in degenerated or herniated discs, suggesting
that degenerated discs are less viscoelastic in nature
(Kazarian, 1972).
Ligaments
Clinical stability of the spine depends primarily on
the soft tissue components, especially in the cervical spine. The spinal ligaments are functional
rnainly in distraction along the line of their fibers.
Ligament strength and limited extensibility help
maintain stability, especially around the craniocervic'll junction. The alar ligaments have an in vitro
strength of 200 N, and the transverse ligaments
have an in vitro strength of 350 N (Dvorak et aI.,
1988a). The strength of the ligaments is related to
both the anatomical demands and the flexibility re~
quired, \vhich is a classic example of form following function.
The ligaments all have high collagen content except for the ligamentum flavurn, which is exceptional
in having a large percentage of elastin. The ligamentum flavum is under tension even when the spine is
in a neutral position or somewhat extended, and th
pre-stresses the disc to some degree and provid
some intrinsic support to the spine (Nachemson
Evans, 1968; Rolandcl~ 1966). The elastic properti
also assist in limiting the in\varcl buckling of the
ligaments during extension, which could potential
compress the neural clements.
Muscle
Muscular strength and control is imperative
maintain head and neck balance. In the cervic
spine, muscle strength also has a role in reduci
stresses on bones. Bending n10ments are applied
the vertebral bodies during various motions. Duri
flexion, tensile stresses are applied to the posteri
cortex and compression to the anterior cortex of t
vertebral body. Substantial loads on the eel-vie
spine have been calculated during neck flexion, pa
ticularly! in the lower cenrical motion segments.
Harms-Ringdahl (1986) calculated the bendi
moments generated around the axes of motion
the atlanta-occipital joint and the C7-TI moti
segment in seven subjects with the neck in five po
tions: full tlexion, slight flexion, neutraL head u
right with the chin tucked in, and full extensio
The load on the junction between the occipital bo
and C I was lowest during extreme extension (ran
ing from an extension moment of 0.4 Nm to a fle
ion moment of 0.3 Nm). It \vas highest during e
treme flexion (0.9 to 1.8 Nm), but this was only
slight increase over that produced when the ne
was in the neutral position. The load on the C7motion segment was low with the neck in the ne
tral position but bccame even lower when the he
\vas held upright with the chin tucked in (rangi
from an extension m0111cnt of 0.8 Nm to a flexi
moment of 0.9 Nm). The load increased somewh
during extreme extension (ranging from 1.1 to 2
Nm) and substantially during slight flexion (reac
ing 3.0 to 6.2 Nm). The greatest loads \vere p
duced during extreme extension, with momen
ranging from 3.7 to 6.5 Nm.
In the same study, surface electrode electromyo
raphy was used to record activity over the erec
spinae muscles of the cervical spine with the neck
the same five positions described above. Intere
ingl}', the values obtained showed very low levels
muscle activity for all positions, even during
treme flexion in which the flexion moment on t
C7-Tl motion segment increased n10re than thr
fold over the neutral position. The fact that the e1
tromyographic levels over the neck extensors were
low in this and other studies (Founlain et aI., 1966;
Takebe et aI., 1974) suggests that the nexing moment is balanced by passive connective tissue SlILICtures, such as the joint capsules and ligaments. This
phenomenon is seen in many other joints in which
passive support is provided by the ligaments.
The values for the moments computed by
Harms-Ringdahl (I986), howcvcr, are appmximately 10 % of the maximal values measured by
lvloroney and Schullz (1985) in 14 malc subjccts
who resisted maximal and submaximal loads
against the head while in an upright siuing position. The mean maximal voluntary moments were
10 Nm during axial rotation or the cervical spine,
12 to 14 Nm during flexion and lateral bending, and
30 1m during extension. Calculation of the maximum (compressive) reaction forces on the C4-C5
motion segment ranged from 500 to 700 N during
flexion, rotation, and lateral bending. and rose La
1,100 N during extension. AnteroposLerior and laleral shear forces reached 260 Nand lION, respectively. Calculated moments and forces generally
correlated well with mean~ll1easured m:voelectric
activities at eight sites around the perimeter of the
neck al the C4 level.
Muscles playa cr-iLical role in basic postural
homeostasis, as can be observed in both historical
and present-day clinical settings. In unique observational studies in lhe 1950s of severclv affecled poliomyelitis patients, improvements in respiratory
assistance for those with respiratory paralysis led
to higher survival rates and a large number of patients who sustained complete paralysis of the cervical musculature. Patients with completely nail
cervical spines were unable to support their heads
unless adequate support was provided and actually
remained in bed despilc the good function of their
cxtrcmities (Perry & Nickel, 1959). Similarly, severc
cervical kyphosis occasionally is seen in elderly patients \vho do not have an obvious stntetural ctiol~
ogy when investigated radiologically. Some of these
patients were found to have marked cenrical extensor muscle weakness that has been attributed to senile cervical myopathy (Simmons & Bradley, 1988)
(Fig. 11-12).
Neural Elements
Biomechanics of the neural elements have noL
been as well studied as the biomechanics of the osteoligamentous vertebral column, but our know!-
lateral cervical radiograph of a 68-year·old 1Noma
presented with severe torticollis. She denied any h
injury and did not have evidence of a suuctural ve
abnormality, infection. tumor, or inflammatory dis
Pseudosubluxations (arrows) are evident subaxiall
consequence of the marked kyphosis. Her neck wa
twisted only because she had a severe cervical flex
formity. and she was unable to see forward excep
turning her head to one side. She was neurologica
tact. It was possible to extend her neck to a relativ
tral position using gentle traction. Following post
sion from (2 to (7, she returned to a normal inde
life. Muscle biopsy was consistent with senile myo
Reprinred with permiSSion from tvioskovlCh. R. (1997).
mSiabihw (rheumarold. d\·varfism, degeneraCtV£!. Oihers
Bridwell &- RL DeW,11d fEds), The Te;(lbook of Spln,ll S
(D.o. 969- i009). Phdacie!pfJJd. Lippincorr-R,wfdl Publishe
edge base is growing. To date, certain bas
meters have been established. The cervica
undergoes significant changes in length
f1cxion and extension (Breig el "I" 1966
1960). Thus, while there is some longitudin
ticity to the spinal corel. it tolerates axial
tion poorl~". It is the translator~' forces llw
cally result in neurological in.illr~·. A comp
tolerance betwecn 2.75 and 3.44 kN is cs
for the adult c~rvical spine before significa
rological injury occurs (i'Vlyers &. \Vink
1995).
Spinal cord injuries can also result from extreme
or sudden flexion-c:\tcnsion movernCl1ts, especially
in the face of a $hallow spinal canal. I-lead Oexion
alone has been shown to result in significant increases in the illlramcdullary spinal cord pressure
in dogs (Kitahara et aI., 1995). Neurological injuries
result from anteroposterior compression of the
,i",··· ..
corel and arc I110re common if the spinal
is stenotic. Flexion motions can result in inwhen the spinal cord makes contact with cervical osteophytes, and cxtcnsion mOl ions may resull
in a pincer-like compression of the cord between
(anterior) ostcophytes and (posterior) invaginated
ligamentum Oavum. Anterior or central spinal cord
injuries may ensue.
Although a diagnosis of spinal stenosis may be
made based on the absolute size of the spinal canaL
imaging of the neuraxis itself may be of greater
value. Contrast-enhanced computerized tomography, myelography, and magnetic resonance imaging
can demonstrate actual impingement or distortion
of the spinal corel. Studies performed in flexion and
extension may enhance the value of the information
by demonstrating the contribution of any dynamic
soft tissue component to the impingement. The exact size of the bony cervical spinal canal and the
vertebral body \vas measured in 368 cadaveric adult
male vertebrae (Moskovich el aI., 1996). This sludy
used well-validated parametric statistical methods
to determine that the mean sagittal diameter of the
spinal canal for C3-C7 was close to 14 mm (14.07 ±
1.63 mm; N ~ 272) (Figs. 11-13 and 11-14). The
mean ratio of the sagittal canal diameter to the vertebral body diameter (canal to body [cob] ralio) was
86.68 ± 13.70. Thirly-one percent of subaxial vertebrae would be diagnosed as having spinal stenosis if
a c-b ratio of less than 80% was considered abnormal. Another study also found a high false-positive
error rate for the e-b ratio, with 49% of 80 asymptomatic football players having a c-b ratio of less than
80O,b at one or morc cervical levels (Herzog et aI.,
1991). Yet another group evaluated the reliability of
the c-b ratio using plain lateral radiographs and CT
scans (Blackley el al., 1999). Results confirmed lhat
a poor con-elation exists between the true diameter
of the canal and the ratio of its sagittal diameter to
that of the vertebral body. The variab.ility in anatomical morphology means that the use of ratios h'om
anatomical measurements within the cervical spine
is not reliable in delermining the true diameter of
the cervical canal.
Spinal cord injul)' without radiographic abnormalities (SCIWORA) have also been described,
N
9i
.
...l2
::"':3
"
C1
••
3.
Oa
N
,
~. a_
I': 17 20 23 26 29
mm
"e
~
'-
.
10 12 14 16 18 20 mm
N
e.
4;
0.__
10 12 14 16 18 20 mm
N 16:
".Il
C4
e·
,.
____ ..
10 12 1': 16 18 20 mm
O~.
N
1.
':.a.
C5
.,
0..
.
10 12 14161820
N 16
".&..
rnfn
C6
e
.;.
0..
.___
10 12 14 16 18 20 mm
N
: ; ...t..-.
C7
81
4'
ot.
_
10 12 14 16 18 20 mm
Histograms of the C1-(7 sagittal canal diameters. Apart
from the (1 plot, the same scale is used on all of the horizontal axes so that the distributions of the diameters can
be compared, Reprinted with permission (rom Moskovich, R..
et at. (996). Does rhe cervical canal [0 body rario predict spina
stenosis? Bull Hasp Joint Dis, 55, 61-71.
especially in children (Dickman el aI., 1991; Osen
bach & Menezes, 1989: Pang & Pollack, 1989). Th
etiology of such injuries is unknown, but one mech
anism may be longitudinal traction. The unusuall
elastic biomechanics of the pediatric bony spine a
lows deformation of the l11usculoskeletal structure
beyond physiological extremes, perm·itting direc
cord trauma followed by spontaneous reduction o
the bony spine (Kriss & Kriss, 1996). The isolate
spinal cord resists tension poorly; axial tensil
rotation or translation of a body ~l!ong one a.\
consistently' assclciated with a simultaneous
tion or translation ~dong another axis, the mot
are coupled. Coupled motions are normally'
pressed as displacements in the X, Y, or Z d
tions and rotations about the three orthog
axes. Further analysis involving whole spine
more complex, and while it has y'iclcled interes
results, motion segment analysis remains im
tant for basic understanding.
Posterior
longitudinal
ligament
Superior facet
vertebra, which was not part of the study detailed in the
Spinous process
text. The anteroposterior diameter of the spinal canal
measured 13.96 mm in this specimen.
•
forces to failure For three adult spinal cord specimens arc reported to be 278 N ± 90 (Yoganandan et
aI., 1996b). Lower forces may result in direct neural
injury or vascular disruption.
Kinematics
Kinematics is the studv of motion of rigid bodies
without taking into consideration other relevant
forces. Kinematics of the spine descrihes the physiological and pathological motions that occur in the
various spinal units. The traditional unit of stud)! in
kinematics is the motion segment, or the functional
spinal unit. As described earlier, each motion segment consists of two adjacent vertebrae and their
intervening soft tissues (Fig. 11-15).
Basic biomechanical testing involves the application of forces to a vertebral body and the subsequent measurement of the movemcnts that occur
(Fig. 11-16). Movements can be either rotational or
translational. A degree of freedom is def-lned as a
motion in \vhich a rigid body can cither translate
back and forth along a straight line or rotate
around a particular axis. Thus, each vertebral body
may either translate or rotate in each of three orthogonal planes, for a total of six degrees of freedom (Panjabi et aI., 1981) (Fig. 11-17). When either
l\v.-o-od{L
Capsular ligament,
Axial computerized tomographic scan of a sixth cervical
"'"
Interspinous
and,
supraspinous
ligaments
II
Anterio
. / longitud
ligame
Interverte
disc
J
I
I p
.
.
ostenor
L'
fl
Inferior facet J longitudinal Intertransv
I ligament
Igamentum avum
ligamen
//
//
I
A
Superior facet I
I
Capsular ligament
Anterio
longitu
ligame
Spinal canal "- '
"Spinous process
~
"' Interverte
disc
Interspinous ~
and supraspinous
ligaments
""_~/
Ligamentum
flavum
~Ls_te_r_io_r
Interlransv
ligamen
Transv
forame
J
Posterior
'!Iongitudinalligament
I
A_n_t_e_ri_o_r
Schematic representations of a cervical motion segmen
composed of two typical cervical vertebrae ((4 and (5)
intervertebral disc, and surrounding ligaments. The bro
line divides the motion segment into anterior and pos
components. A, Lateral view. B, Superior view. Adapied
permission from White, A.A. 1/1, Johnson, fUvr, Panjabi, lvUv
al. (7975) Biomechanical analysis of clmica! stability in rhe c
cal spifJ(l. (lin Orthop. \09,85··-96
n
, ,
9
:
i
I
c..;
Diagram of a test rig to evaluate a functional spinal unit
using videophotogrammetry. This technique facilitates accurate measurements of motion without the measurements themselves having an effect on the displacements of
the mobile vertebra. A. light-emitting diodes (LEOs); Band
C. Guide bars for application of tensile and compressive
fOfces. D. Pulley for application of torques. Weights are at·
tached to guide wires (w), which go around pulleys to produce torque on the upper vertebral body (v). n. intervertebral disc; 9. acrylic cement that attaches lower aluminum
plate (m) to test rig, which is rigidly bolted to the support
frame. The upper plate (p) and upper vertebral body (v)
are the moveable elements to which the loads and torques
are applied. LEOs are rigidly attached to the upper plate
and their movement is recorded by two video cameras.
;·Fy
c1
Reprinted with permission {rom Raynor. R.• Moskovich. R.• Zide/,
P. et a/. (1987). Alteration in primary and coupled neck motio1J5
l__
I
m
aled is the fact that a considerable amount of fle:x
ion and extension occurs at the CI-C2 articulation
5 to 20° or flcxion and extcnsion occur with mean
of 12" (aetive) to 15° (passive) (Dvorak et aI., 1988b)
Approximately 90° of axial rotation takes place
in the subaxial eel"vic'll spine (C3-C7), about 45° to
cach side of ncutral. Even greater lateral flexion i
possible: approximately 49° to each side of ncutral
giving a total of about 98°. The range of flexion and
extension is approximately 64°: about 24° of cxten
sion and 40° of flexion. The motion in each plane i
fairly evcnly distributed lhroughollt the motion
segments (Lysel!, 1969). Thc mean total range o
anteroposterior translation in subaxial spinal mo
tion segments is 3.5 ± 0.3 mm divided unequally
1.9 mm for anterior shear and 1.6 111m for posterio
shear. Lateral shear loading results in a mean tota
range of lateral motion of 3.0 mOl :!:: 0.3 mm, di
vided equally between right and left; tension result
in 1.1 mm of distraction and compression, 0.7 mIl
of loss of vertical height (Panjabi et aI., 1986).
The great Ilexibility of the eel'vical spine allow
the head to be positioned in a wide variety of ways
a_f_,e_'_fa_,_e_,e_,_,o_n_'Y_._N_,e_"_,o_S_"'_9_e_ry_,_2_1(_5_),_6_8_'_-_68_7_.
_
l1Y
RANGE OF MOTION
Measurements of cervical range or motion are based
on radiographic studies or postmortem investigations. Inclinometers and various optoelectronic and
electromagnctic dcvices L1sed clinically for noninvasive evaluation of cervical spinc motion are not as
accurate; in particular, coupled motion is poorly
quantified (Roozmon et aI., 1993). The established
range of active axial rotation to one side at CI-C2 is
27 to 49° (mean = 39°); passive rotation is 29 to 46°
(mean = 41°) (Dvorak et aI., 1987; Dvorak et aI.,
1988b; Penning & Wilmink, 1987). These measurements account for approximately 50 r1'o of the lotal
cervical rotation.
Another stereoradiographic study of neck motion
in adult men found a mcan of 105° axial rotation between the occiput and the C7 vertebra. Seventy percent of the total axial rotation OCCUlTed between the
occiput and the C2 vertcbra. Eaeh motion segment
between the C2 and C7 vertcbrae avcraged fTom 4 to
8° rotation (Mimura et aI., 1989). Less well appreci-
+Fz
~
0i---le.
~Fx
A vertebral body showing the three primary Cartesian
axes, x, y. and z. Along each axis a positive force, ~ F. is de
noted by the direction of the arrow. The curved arrows indicate the direction of a positive torque. +..... Reprinted will
permission from Raynor, R.• Moskovich.
R., Zidel, P., el af.
(/987). Alteration in primary and coupled neck motions aller
facete<:tomy. Neurosurgery. 21 (5). 681-687.
;
i
l-",~~=~,...,.-.,-,~~-=-=-=--~;==;r.----,.--------:=,,~,
'.""'" .-.&f
""
..
",-"_",,,,,""£1 -
.~"""'"- .-~-
~.~.
•__ ' -,
permitting one, with equal ease, to gaze at an air~
plane overhead, glance over one's sho111dcI~ or look
for an object under a table. An analysis of the com~
bined motion of the cervical spine Llsing an eleclrogoniometer produced a remarkably large range of
motion (Feipel et aI" 1999): 122' ;;; 18' of nexion and
extension, 144°:!: 20 0 of axial rotation, and 88 0 ::': 16°
of lateral nexion, All primal)' motions were reduced
with age. Sex had no influence on cervical motion
range.
The active range of cervical spine motion required to perform daily functional tasks was studied
in healthv adults (Bennett et aI" 2000), A cervical
spine range of motion device was fastened to the
subject's head with a Velcro strap and a magnet was
placed on the patient's shoulders to calibrate the instrument for measurement of cervical rotation and
motion, Of the 13 daily functional tasks performed,
tying shoes (flexion-extension 66.7°), backing up a
car (rotation 67.6°), washing hair in the shower
(flexion-extension 42.9°), and crossing the street (1'0·
tat ion head left 31.7° and rotation he4:KI riglll 54.3°)
required the greatest active range of motion of the
cervical spine. Of interest, several tasks were not
found to produce the degrees of lnotion expected,
and these included reading a newspaper (19.9°
Ocxion-cxtcnsion), writing at a table (26.2° flexionextension), and reaching 1'01' objects overhead (4.3°
flcxion·extension). Side·bending was not found to
be a significant movement in completion of the
tasks but was coupled \vith rotation in one of the
tasks (looking left and right to cross a street).
SURFACE JOINT MOTION
The motion between the joint surfaces of two adjacent vertebrae Illay be analyzed by means of the in·
stant cenlct· techniquc of Rculeaux, described in
Chapter 6. The method may be used to analyze sur·
face motion of the cen:ical spine during flexion·
extension and lateral flexion.
In a normal cel'vical spine, the instant center of
flexion-ex lens ion is located in the anterior part of
the lower vertebra in each motion segment. Instant
center analysis indicates that tangential motion
(gliding) takes place between the Facet joints as the
cel"ical spine is flexed and eXlended (Fig, 11-18), A
consequence of these rnotions is that the size of the
intervertebral foramina increases with flexion and
decreases wilh extension (Fielding, 1957), These alterations have been quantified in a cadaver study
that found there were statistically significant reductions of 10 and 13% in foraminal diameter, at 20 and
\
!
II
C5
Analysis of surface motion of the facet joints of the
motion segment during flexion-extension. The sche
drawing represents superimposed roentgenograms
motion segment in the neutral position and in sligh
ion. The upper vertebra ((4) is considered to be the
ing body and the subjacent vertebra ((5) is the bas
bra. Two points have been identified and marked o
moving body in the neutral position (sofid outline o
and the same two points have been marked on the
roentgenogram with the motion segment slightly f
(dashed outline of e4). li"es connecting the twO se
points have been drawn and their perpendicular bi
have been added. The intersection of the perpendi
sectors identifies the instant center of motion (larg
dot) for the degree of flexion under study. The per
ular bisector (arrowed fine) of a line drawn from th
ter of motion to the contact point of the facet join
faces indicates tangential motion. or gliding.
.-,-~~~~~--~-~~-
30° of extension, r('spccti\'c1~'. Convl'rsel~", in
lh~r~ \\'cre statistically significant increases o
10% at 20 and 30° 01" fll.:'xion, l"l.'spcclin~ly (Yoo
1992), One p""clical application of this dala
to cervical collars lIsed for the rdief of neck
Conventional foam collars tend to place pati
slight extension, which IlU\.\' aggravatc the
toms. Turning a foam collar around. with thl.:'
and narrow parl anterior, puts (he neck in
flexion, which may increase the size of the in
tcbral foramina and thereb~1 relieve some
pressure on an inflamed nerve root.
The instant center
motion of' the cClvica
may be displaced as a result of patho
processes such as disc d~gcncrati()n or ligam
pairment. In such cases, instant center analys
reveal distraction and jamming (compress
the facel joint surfaces during flexion·extens
sl~ad or gliding (Fig, 11-19),
or
C4
AnI.
Post.
C5
I
Schematic drawing representative of a roentgenogram of
the C4-CS motion segment of a patient injured in a rearend auto collision. The instant center of flexion·extension
at this level (represented by the large solid dot) has been
displaced from the anterior to the posterior part of (5 as a
result of the injury process, which impaired the ligaments
(compare with Fig. 11·18). The analysis of surface motion
shows compression and distraction of the facet joints with
flexion and extension.
•
there are 2° of coupled axial rotation for every 3
lateral bending, which gives a ratio of 2:3 or 0".67
C7, there is lOaf coupled axial rotation for ev
7.5 0 of lateral bending, which gives a ratio of 2: 1
0.13 (White & Panjabi, 1990). Results of finite
ment rnodeling indicate that the facet joints and
covenebral joints are the major contributOl"S to c
pled motion in the lower cervical spine and thal
uncinate processes effectively reduce motion c
pling and primal)1 cer'vical motion (in the same
rection as load application), especially in respo
to axial roLation and lateral bending loads.
covcrtcbral joints appear to increase primal)' ce
cal motion, showing an effect on cervical mot
opposite to that or the uncinate processes (Clau
et aI., 1997). Coupling or nexion-exLension w
transverse translation may be visualized roentge
graphically (Fig. 11-23). During flexion, the VC
bral body normally shifts rorward: the racets g
up and over one another with pseudosubluxali
Changes in normal coupling patterns occur foll
iog palhological changes or surgical intervention
Tensile load application to the cervical sp
occurs in therapeutic traction, but more co
monly so during trauma. Deployment of pass
COUPLED MOTION OF THE CERVICAL SPINE
Atlantoaxial Segment
i:
The coupling characteristics of the atlantoaxial
spinal motion segments arc particularly important,
as this area of the neck is extremely' mobile. The
dens is constrained within the osteoligamentous
ring of the atlas, causing the C l-C2 lateral masses
to articulate similarly to the condyles of the knee,
with some sliding and rolling during Oexion and extension. The instant centers of both rotation and
n~xion-extension lie in the center of the dens itselr.
Rotation at CI-C2 is coupled with both vertical
translation along the Y axis (Fig. 11-20) and a degree of anteroposterior disphlcenlent (\,Verne, 1957).
This implies that the CI-C2 joint is most stable in
the neutral position, and, if rotated, allempts should
be made to return it to the reduced position when
performing an arthrodesis.
Subaxial Spine
The coupling patterns in the lower cervical spine arc
slIch that in lateral bending to the left, the spinous
processes move to the right, and in lateral bending
to the right. they move to the left (Lysell, 1969:
!vloroney et at.. 1988) (Figs. I 1-21 and 11-22). AL C2 .
.
_----~"
[
~'
\
\
A
\1'
B
Coupling of rotation and axial translation is depicted
schematically. A, (1 and (2 are in the neutral position.
S, C1 rises fractionally on C2 (arrow) as the head is rota
away from the midline. Adapted with permission [rom
of
Fielding, J.W (1957). Cineroentgenography
the normal
cervical spine. J Bone Joint Surg. 39A, 1280-/288.
-:__
I'
vehicular n:.'S!l"airH s.\'stCtllS, such as airbags,
induce tensile forces in the neck. Isolatcd
vertebral discs fail at 569 N
54. and intac
man cadaver cervical spines rail at 3373 N ,
(Yoganandan cl aI., 1096h1. Active muscular
traction, 110\\'C\'i..'r. is liki.'I~' 10 raise these fi
cons i dl'J"~1 hi v.
Abnormal Kinematics
Left lateral
flexion
Neutral
Righi lateral
flexion
Abnonnal kinematics gCl1crall.\" refers to l'.'\c
motion within rUl1etional spinal units; howeve
norrnal kinClnatics Ina.. . · also refer to at.\vica
Coupled motion during lateral bending is depicted
schematically. When the head and neck are flexed to the
left, the spinous processes shift to the right, indicating rotation. The converse is also illustrated. Adapted with permission from White, AA 1// (~Panjabi, MM (1990). Clinical Biomechanics of the Spine. Philadelphia: 1.8. Lippincott.
A
A, This diagram illustrates some of the other coupled motions, which occur in response to a torque (liZ) about the Zaxis (lateral bending). lateral translation (Rx) and vertical
motion (Ry) occur, as well as horizontal rotation (l!lY), which
terns of motion such as abnormal coupling or
doxicalmotion. Par~\do.\icallllotionis seen whe
overall pattern of motion of one ~lSPCCI of the
is in one direction while the local pattern is th
posite. For instancc, parado'\ical llc'\iol1 is
when llc'\iol1 occurs at a single functional s
unit, although the SpillL' as a whole is (''\te
B
results in motion of the spinous processes to the right
left. B, The subject is bending to her right, demonstrat
the large range of normal cervical motion possible (ap
mately 50°).
I
I
Coupling of flexion-extension with transverse translation of
the cervical spine is visible roentgenographically. A, Ouring
flexion the vertebral body shihs forward (small white arrow); the facets glide up and over one another with moderate subluxation at full flexion (large white arrow). Up to 2.5
mm of transverse translation may normally occur at the
(1-(2 articulation during flexion-extension; no translati
evident in this example (black arrow). B, During extensio
the reverse occurs, and the spinous processes limit motio
they touch at full extension (arrow). The size of the inte
tebral foramina increases with flexion and decreases wit
extension.
•
These types of abnormal motions describc a pattern
of mOVCIl'lCllt known as instability.
Spinal Stability
The concept of spinal stability is an intriguing and
sometimes confusing notion. Medical practitioners
are frcqucnlly asked to look at a sel'ies of radiographs and make a dctermination whether the
spinc is stnblc. Exnclly what is stability, how is it determincd, "'ld what happens if it is not present? The
term spinal stability has acquired diffcrent mcan~
ings, depending on the setting in which it has been
used. White and Panjabi describe the term clinically
as the loss of the ability of the spine undcr physio-
logical loads to maintain its pattern of displacem
so that there is no initial or additional ncurolog
dcficit, no major defornlity,' and no incapacita
pain (1990). Using this definition, physiolog
loads are those incurred during normal activity,
pain is not felt to be incapacitating if it can be c
trolled by non-narcotic drugs.
Stability is detcrmined by many factors. There
different anatomical considerntions in different
gions or the spine. Certainly, ligamcntous anato
plays a large part in the stability or the spine, but
muscular and bony clements of the ~pine also
important roles.
Instability can be analyzed by considering kinem
instability and structural or component instabi
Kinematic instability focuses on either the quantity of
motion (too much or too little) or the qualit..., of motion
present (alterations in the nOimal pattenl), or both.
Component instability addresses the clinical biomechanical role of the various anatomical components of
the functional spinalunil. In this type of instability, los..o:;
or alteration of various am.llomical portions detennines
the presence or instability (Box II-I).
Sir Frank 'Holdsworth's account of a simple twocolumn concept of spinal stability provided a constructive basis for describing and analyzing the basic biornechanics or the spine:
The synal1hroscs bClwcen Ihe \'cl1cbral bodies rely for
theil' slabilit~, upon the immensely slrong annulus fi~
brosus. ThL: clim'lhrodial apophys(.';:d joints arc slabilised by the capsule, by rhc interspinous and
sllpraspinous Iigamenls and the ligamcnta Ila\"a. This
group of ligaments [ call the 'postedol' ligament complex.' It is upon this complex (hal (he swbilily of lh~
spine largdy dq:>ends (Holdsworth, 1963).
Subsequently, Denis (1983) described a classification system for thoracolumbar fractures that can
also be applied to a biomechanical analysis of spinal
stability. In this description, the spinal elements are
divided into three regions that form three spinal
columns:
I. The anterior column consists of the anterior
longiludinalligament, the anterior annulus fibroslls, and the anterior half of the vertebral
bock
2. The middle column consists of the posterior
longitudinal ligament, the posterior half of
the venebrnt body, and the posterior nnnulus
librosus.
3. The posterior column consists or the pedicles,
facet joints, lam inn, spinaliS processes, as
well as the interspinous and supraspinous ligaments.
4. The anterior and middle columns form the
primary weight-bearing column of the spine,
with the posterior column providing the guiding and stabilizing elements.
Occipitoatlantoaxial Complex
The transverse ligament of the atlas completes the
socket into which the dens is inserted. The ligament allows the dens to rotale, but limits its anterior translalion. The ligament is inelastic and will
not permit more than 2 to 3 mm of anterior subluxation of the first on the second vertebra (Field-
Conceptual Types of Instabilit
_
Kinematic Instability
Motion increased
lm.tantaneous ax(,>s of rotation altered
CDlipiing characteristics changed
Parddo;O:I(d\ mOllon present
Component Instability
i"raun1C1
TUIl10!
Surgery
Degenerative changes
Developmental changes
Combined Instability
K:nemiitK
Cornponen!
ing I..'t aI., 1974). Antcrior displaceml... nt or CIo
or 3 to 5 mm is usuall:' indil';:ltive or a ruptu
till' tr;:\Ils\\:rse ligament. while disphlCl'Il1ClHS o
10 mill suggest nCCl.:'Ssol'\ lig<.\JlK'llt d~l11lage;
plncC'tlll'lll greater than !O lllill occurs with ru
or all lhe ligal1lL'lllS (Fil.'lding et :.II .. 1(76). An
translations or displacemcnts of Cion ('2 :J,
sessed I'adi()graphicall~' by measuring tilt..' dis
from lI1L' anteriur ring of the alias In the back o
dc'J1' (atlaJ1todent,,1 intcly,,1 or ADI) (Fig. 1
(CasL' Stud.'" 11-1). Poslt.:'rior subluxation of lhe
C<.lll onl~' OCCUI" if lhe dens is fractured or if tllL
all OS-OdOlllOidL'ulll or hypoplastic dellS.
SOl11e diseases can wcakc'll or dL'S{nl~' th\.'
n.'l'sc ligament. :\,10S1 nOlahl...-. s.vllo\'itis in rhL
!oid arthritis can crL'~ltc a pannus. which hel
d(;SII'O,\' the atlantoa.\:ial ~1J'liClllalion as well a
ll'ansn'!'SL' Iigamcnt (Figs. 11-25 and 11·26L
tients with Do\\'n s.\'ndroll1c <.l1'C also susceptib
\\,c;:lketlcd tmnsvcrsc ligaments and must be
full." assessed clinicall.\" and radiographicnll .., b
bdng allowed 10 pankipalc in sporting events
as the Special OI~'l1lpics.
Stccrs "Rule of Thirds" is a guidt: 10 the am
or allanloaxial displw':cl11cl1t that can OCCllr b
spinal cord comprcssioll ensues (Ste~1. 1968)
illt~rnal anteropO-SIL'l"inr diameter
pl'o.,\jJl1alel~·
01" thL' atlas
3 cm: of thal. the dens occupie
proximately I cm and thl' spinal cOI"(1 approxim
I cm. kadng I CI11 of space for soh tisSlh.' an
normal mo\"cnwnt to OCClll" (Fig. 11-27).
subaxial Cervical Spine
(1963) stated that the musculature of the
spine and the intervertebral discs were the most
significant anatomical structures is providing cer'vie'll stability. Holdsworth (1963) emphasized the
imparlance or the supraspinous and interspinolls
ligaments as well as the nuchal ligament. The
nuchal ligament is thought to playa major role in
Atlantoaxial Instability Without Fracture
f\:},9;Ye'a.r~old woman had a traumatic injllry as a result
.'co"J;t!.of a forced flexion movement in a car acodent. Conhinuous and severe neck pain occurred after the accident.
(s.,he visited the emergency room and after a careful ex·
~~~~rr,ination and x-ray evaluation, an anterior displacement
~'ZR.f:Cl
on
(2 was discovered (Fig, cs11-1·1).
~:(E:::;A severe amerior dislocation of the atlas on the axis
~~:/~~as
confirmed
/':"__i"::",'"
'.,1'
SAC
after measuring the atlantodental interval
. (~'rnrn). ,~or this case, no fracture of the atlas or the axis
\i\I~~"~:7,:te~,t,~d and thus a defiCiency of the transverse ligalfIent,:'r;na'y, be assumed.
Clinical instability of the spine depends mainly on the
soft tissue components. The cervical spine is
a very mo·
,bile area, especially at the atlanlOaxiallevel. Cervical subluxations and dislocations resulting from injuries of the
osteoligamentous complex affect spinal stability and rna~ility at the cervical area. In addition,
it may narrow the
spin,al canal and cause neurological impairment. Pure lig.
am~ntOtJs injuries
(as in this case) are less likely to heal
and become stable, and therefore surgical procedures are
lik,ely to be considered.
SAC
I
I
~
I
I,
L.-
_
The atlantodental interval (AD!) is inversely related to the
space available for the spinal cord (SAC), which is denoted
by the dotted lines. Anterior atlantoaxial subluxation
causes a reduction in the SAC. Normal measurements for
the ADI are less than 3 mm in adults or 4 mm in children.
Reprinted V'lirh permission from Moskovich, R. (1994). Atlanroaxial instability Spine: Stare of the Art Reviews. 8(3), 531-549.
-rease. Study' Figure 11-1·1. lateral radiograph demonstrating
. ,,~,?j}.n:~,~!~?sed atlantodental interval of 6 mm after trauma.
proprioception and correct functioning of the
erector spinae muscles. Experimental section o
the ligaments in sequence from either anterior to
posterior or posterior to anterior suggests L1lat i
a functional spinal unit has all or its anterior ele
ments plus one additional structure or all or it
posterior elements plus one anterior structure in
Flexion lateral radiograph of a patient with rheumatoid
arthritis. The dens (0) has been eroded and the transverse
ligament is incompetent, resulting in atlantoaxial subluxation. The markedly widened atlanto·dens interval is indicated by the arrowed line.
tact, it will probably remain stable under normal
physiological loads. To provide for some clinical
margin of safety, any malion segment should be
considered unstable in which all of the anterior
clements or all the posterior elements arc destroyed or are unable [0 function (Panjabi eL al..
1975; White et al.. 1975). The clinical sLability of
various injuries must be assessed individually.
The importance of clinical evaluation cannot be
underestimated because significant spinal corel
damage may occur after trauma even in the absence of fractures or ligamentous injuries (Gosch
et al.. J972; Schneider et aI., 1954). Valuable
guidelines ror the detenllination of clinical instability in the lower cervical spine have been provided in the form of a scoring system checklist
(Box 11-2).
Using this scale. the measurement of translation takes into account variations in magnificalions and is based on a tube-ta-film distance of
183 em. The 11 rotation is defined as 11 greater
than the amount of rotation that exists at the motion segment above or below the functional spinal
unit in question. The 3.5-mm value represents the
radio-graphic measurement of the maximum permissible translation when the radiographic magnification is taken into account (Panjabi et aI.,
1986).
0
0
1DmI,
!
,
1 ....d_e_co_m.pression by transoral resenion of the dens c
1
,..
ItM'~'"'~"~~M" .' A~_.~' ,~;;;Sk0.
•.,••. ·W·,·_<· ,,,--
',<>'<
'.
".....
terior stabilization.
III
C~lS(.'
or
Cjl1l...·slicJl1able
in.iury. when
~Hl(1
extension ill~lI)Cll\'l'rS should Ilol b
formed, a stretch Lest l'~lll b~ d01l1:.' 10 asse
deal inlq.!rit~·. A la(I.'ral CI.·ITiGll spine
ograph is taken al a standardized lube d
of 180 em. ~\lld incrclll(,.'nud .i-kg weights
plied as traction \n lhe Sklllilisillg LT<.lllia
(Fig. 11-28). RadiogT~\phs arc tak~n ark
addilion~d
/'i,'
hl~~
iPi
Postmyelogram cervical radiograph in a patient wi
standing rheumatoid arthriris and fixed atlantoaxi
luxation. The space available for the cord has been
duced to 6.5 mm and compression of the proximal
cord is evident by examining the subdural space. w
outlined in white by the injected contrast medium
tient developed cervical myelopathy. which necess
.''''''~'
wcight has been npplicd. An
Dens
1
"3
Space available for the cord (SAC) is approximately twothirds of the anteroposterior diameter of the spinal canal.
One-third is taken up by the dens, one-third by the cord itself, and one-third is free space. Reprinted with permission
;rom MoskDvich, R. (994). AtlanrD-axial in5tclbility- Spine: State
Df (he Art Reviews. 8(3). 531-5/19.
Stretch test. A, Lateral radiograph of the cervical spine o
19-year-old boy who was admitted with multiple injuries
and neurological deficit compatible with an anterior cor
syndrome. The radiograph shows increased angulation a
C3-C4 and a block vertebra at (5-(6. Flexion-extension r
diographs were contraindicated for fear of exacerbating
his neurological injury. B. The stretch test was performed
to ascertain whether significant instability existed. No ab
normal distraction occurred at the interspace in question
His other injuries and fractures were treated routinely an
he was given a soh collar to wear for 6 weeks. He made
slow but steady recovery with almost complete resolutio
of his upper extremity deficit. with no evidence of instab
ity 1 year after the accident. Reprinted with permission (rom
Moskovich, R. (1997). Cervical insrability (rheumatoid, c!warfism
degenerative, others). In K.H. Bridwell & R.i. DeWald (Eds.), T
Textbook of Spinal Surgery (pp. 969-1009). Philadelphia: Lipp
•
mal Slretch tesl is defined as differences of
·'greater than 1.7 mm 'interspace separation or
greater than 7.5 change in angle between prestretch condition and the application of Ot1Cthird body weight.
0
•
I
I
I
Checklist for the Diagnosis
of Clinical Instability in the
Middle and Lower Cervical Spine
cott-Raven Publishers.
Point
Value-
Element
•
,A,nteri~r elements destroyed or unable to function
I fJostenor elements destroye or unable to function
I Positive strelch test
I
!
!
Applied Biomechan.ics
4,
Radiographic criteria
A thorough understanding of biomcchanical princ
Flexion and extension x-rays
ples is an important aspect of the treating phy
cian's knowledge base because the normal structu
and function of the spinal ,column is frequent
altered during surgcly. \'''hether treatment is a d
c:omprcssive cervical laminectomy, a posteri
I-oraminotomy with partial facetcctolllv, or an ant
rior cervical fusion, all of these inlen~'cnlions ha
certain ramifications with which onc must be fmn
iar. This kno\vledgc not only benefits patient care b
also is valuable in planning and executing treatmen
Sagittal plane translation> 3.5 mm or 20% (2 pts)
I
Saginal plane rotatiDn > 20° (2 pts)
OR
I
Resting x-rays
Sagiltal plane displacement> 3.5 mm or 20%. (2: pts)
Relative sagittal plane
angulation> 11° (2 pts)
I Developmentally narrow spinal canal
I
.:\bnorrnal dISC narrowing
I Spinal cord damage
2'
!
1
Nerve root damage
I Dangerous loading anticipated
1
DECOMPRESSION
-"' Total of 5 or more == clinically unstable.
I
blodifiecJ from While. A.A. III & Panjabi. M.M. (1990). Clinical Biome<hdnrcs oi the Spine (2nd ed., p. 314). Philadelph;ej: 1.B. lippincott.
•
Cervicallaminectom)' often is performed to decom
press the spinal corel. The compression may
caused by a stenotic process and can result
neurological sy'mptoIlls such as racliculopath.v or
my'clopath)'. Other posterior decompressive procedures such as partial or full facctcctomics for visualization or decompression of nerve root pathology
also arc comnlonly performed. Development of
postlaminectol11)' kyphosis is well known in children and may develop in 17 (Miyazaki et a!., 1989)
to 25% of adults (Berkowitz, 1988). [,ostlaminectomy spinal deformity may..' occur in up to 50% of
children who undergo laminectomies for spinal
cord tumors (Lonstein, 1977). Simulated finite element anal.'.'sis on cervical spines indicates that the
primary cause of postlaminecto!l1 yr deformity is resection of one or more spinolls processes as well as
the posterior ligamentous structures, such as the
ligamentum flavulll or interspinous or supraspinous
ligaments. The removal of these structures causes
the tensile forces normally present in the cervical
spine to become unbalanced and place extra stress
on the facet joints. Results indicate that either a
kyphotic or a hy'periordotic cervical defonnity' may'
ensue, depending on the center of balance of the
head (Saito et aI., 1991).
Although reports exist of patients undergoing
multiple-level cervical laminectomy with no evidence of clinical instability or deformity' on longterm follow-up (Jenkins, 1973), most biomechanical studies indicate some degree of instability when
the posterior elements arc resected. Multilevel cervical lan1inectom.v induces significant increases in
total column flexibility associated with increased
segmental flexural sagittal rotations. In a cadaveric
laminectomy model. the mean stiffness of the intact ce,·vical column was significantly' greater than
the mean stiffness for the laminectomized specimen, and there were consistently greater rotations
as compared with the intact specimen (3.6 versus
8.0°) at ever)' cendcal spine level (Cusick et al.,
1995).
The loss of facet joints alone causes a significant
decrease in coupled motions that result from lateral
bending. A moment about the anteroposterior axis
results in a significant reduction in lateral displacement, a decrease in vertical displacement, and a decrease in rotation about the vertical axis. Partial
facetectomy «50 % ) did not, however, significantly
alter flexion and extension movements (Raynor et
a!., 1987). Another anatomical study demonstrated
that progressive laminectomy with resection of
more than 25% of the facet joints resulted in signif~
icantly increased cervical flexion-extension, axial
torsion, and lateral bending motion when compared
\vith the intact spine (Nowinski et a!., 1993).
A nUlllber of studies using three-dimens
nite element models showed that f::\Cetectom
greatcr errc'ct on annulus stress th<:ll1 (Hl in
bral joint stillness. Based on these models
concluded that a signific<:lnt increase in
strcsses and scgmentalrnobility nlaY ()ccur
lateral facet resection c:\:ceeds S(Y'·'r (Kuma
a!., 1997; VClO et a!., 1997), Deccnnpression
cal laminoplasty, in which the facet joints
sacrificed and the laminae are reconstructed
in maintenance ()f fle:\:ic)n-e:\:tension and
bending stability', \vith a nwrginal increase
torsion. Iatrogenic injur.v is less likel:' to res
capsules of the rernaining facet joints and
elements remain intact.
Cervical subluxations and dislocations
from injul':' may narrow the spinal canal an
neurological imp::lirmenl. In SOllle cases, a
reduction and re::dignment or the verteb
lowed b~' stabilizatkm, may' decompress th
elements without resecting bone (Fig. 11-29
ARTHRODESIS
Spin::d arthrodesis is indicated in man:' dise
cesses such as spinal instability, neoplas
traumatic and degcnerative conditions of th
Thc goa! of arthrodesis is to achicve a solid
sion in which two or more vcrtebrac solidi
In 1l1an:' cases, internal fixation is ust::.'d to
initial st::\bilization as well as to correct defc
necessary.
An important principlc regarding v
arthrodesis is that the stabilit.\' established
nal fixation is a prelude to the biological p
fusion. Although fusic)l\ can occur without
fi:\ation, its appropriate USl' helps to increas
sion rate and maintain structur::ll alignmen
ware b:' no means supplants the necd for
gecm to perfol'ln a thorough and careful pre
of the vertebrae and add bone graft tCl ach
sion. \Vith few exceptions, Imrdw::lre that is
matel.v supported and protected by a soli
will fatigue and fail after a finite number o
There is ~l race b." the body to achieve a sol
mass before a fatigue failure (If the implan
tion device occurs,
The choice of a surgical approach !() the
spine, as well as whether an anterior, poste
combined arthrodesis should be performe
pendent on the particular patholog." presen
performing a fusion, it is important for the
to understand the biomeclwnical propertie
Unilateral facet dislocation in a 24-year-old woman who was involved in a motor vehicle
accident. The degree of vertebral subluxation is less than half the AP diameter of tbe
vertebral body. Her spinal cord was compressed and she presented with an incomplete
spinal cord injury. A, Computerized tomographic scans and reformatted images demonstrate the canal compromise at (5-(6. An attempt to realign the spine was made using
longitudinal traction of up to approximately one-third of her body weight. B, The lateral
radiograph with traction applied shows persistent malalignment. An open reduction was
performed using a posterior exposure of the vertebra. Once the spine was realigned, the
posterior tension band was recreated using interspinous wiring, and autogenous bone
graft was inserted. C, The radiograph taken after the operation demonstrates restoration of normal vertebral relationships. The fixation provided good stability and enabled
the patient to mobilize early. She had a good clinical outcome. Reprinced with permission
from Mos,l:ovich. R. (1997). Cervical instability (rheumatoid. dwarfism, degenerative. others). In
K.H. Bridwell & R.L. DeWc1/d (Eds.). The Textbook of Spinal Surgery (pp. 969-1009). Philadelphia:
Lippincott-Raven Publishers.
fcrent
t)'PCS
of fusion constructs. Cervical arthrode-
sis also has an affect on adjacent motion segments.
Thcorcticall.',/, there is an increase in motion at
nearby' unfused levels. Subsequent degeneration of
other motion segments requiring additional levels
of arthrodesis has been demonstrated in multiple
studies (Cherubinc) et aI., 1990; Hunter et aI., 1980).
A recent study by Fuller ct al. (1998) evaluated the
distribution of motion across unfused cervical motion segments after a simulated segmental arthrodesis in cadaver cenrical spines. The authors simulated
01112-, t\\'o-, and three-level fusions in human cervical
spines. They then moved the cervical spines through
a nondestructive 30° sagittal range of motion and
compared this range of motion with that of unfused
cervical spines. The findings of this study were interesting in that sagittal plane rotation was not increased dispropc)J'tionatel)! at the cervical motion
segments immediately adjacent to a segmental
arthrodesis. Although the authors ackno\vledged
certain limitations of the study, they proposed that a
cenrical fusion causes a fairly.' uniform increase in
motion across all remaining open cervical motion
segments; therefore, an increased potential for degenerative change may exist at all cervical levels.
Another stud.y was performed descdbing the incidence, prevalence. and radiographic progression of
symptomatic adjacent level disease after cervical
arthrodesis (Hilibrand et aI., 1999). Adjacent level
disease \vas defined as the development of a new
radiculopath,Y or myelopathy that was referable to a
motion segment adjacent to the site of a prior anterior cervical arthrodesis. The findings revealed that
symptomatic adjacent level disease occurred at a
relatively constant incidence of 2.9 % per yeac A survivorship analysis revealed that approximately 26(Y£)
of patients \vha had an anterior cervical arthrodesis
would have new disease at an adjacent Icvel within
lO y"cars of the operation. The study also demonstrated that more than two-thirds of the patients
who developed adjacent level cervical disease experienced failure of nonoperative treatment and
needed an additional procedure performed.
tems have become available that can satis
stabilize the cervicul spine from either app
\Vhen an anterior approach is used, a dis
or a vertebrectOlny at one or more levels is co
performed, usuaIlv
- followed bv. an anterior
arthrodesis. The excised disc or bone mu
placed with a structural graft or prosthesis t
anterior column support to the spine. Oss
placement ma:v be in the fm'm of autogenou
genic bone, comrnonly from the iliac crest i
nous or from t.he fibula or iliac crest if alloge
11-30). The more cortical nature or flbula gr
result in dela.yed incorporation compared"" w
crest grafts, which, therefore, are preferred c
The immediate postoperative strength of an.v
bone grafts under axial compression on a
testing machine reveals that they \\'ill adequa
port the loads required in the cervical sp
ph.\'sical properlies of human bank bone se
preferable to autologous bone grafts, espe
Cervical Spine Fixation
Arthrodesis of the cervical spine Illay be indicated
for a number of reasons, most commonly for
trauma and degenerative diseases. lVluch research
hus been performed to analyze the biomechanical
advantages of anterior approaches, posterior approaches, or a combined procedure. "Vith the advent of newer technologies, internal fixation s.\'s-
Lateral radiograph of the cervical spine of a 35-y
male who had an anterior cervical discectomy at
interbody arthrodesis using autogenous iliac cres
bone graft. Note the maintenance of lordosis at
segment and the integration and remodeling of
Compressive Strength of Various Interbody Graft Materials
Graft Type
Mean ~
Standard Deviation
Fibular strut
5.070" 3.250 N
Anterior iliac crest
1.150" 487 N
Posterior iliac crest
667,,311N
Rib
452 " 192 N
Hydroxylapatite
200 f-l.m pore size
1,420" 480 N
Hydroxylapatite
500 f.lm pore size
338" 78 N
Fibular strut significantly
stronger than crest or rib
grafts: P < 0.05
Pore size of 200 ~m signifj·
cantly stronger [han 500 j-Lm
pore size: P < 0.05
Adapted with permission itQm Wittenberg, R.H.. Moeller. J., P. White, AA Ill. (1990). Compressive strenglh
of autogenous (lnd alfogenolls bone grafts for thOf.lcolurnb.lf and cervical spine fusion. Spine. 15(0),
1073-1077.
older patients (Wittenberg et aI., 1990). The biological
incOIvonuion or allograft appears to be satisfactory
and its USe obviates the need to harvest autologous
bone from the iliac crest, sparing patients additional
surgical trauma and potential complications.
Calcium phosphate (bone) ceramics form a strong
bond with host bone because of a zone of apatite microcrystals deposited perpendicular to the hydroxylapatite ceramic surface (Jarcho, 1981). Synthetic
hydroxylapatite blocks used for interbody cervical
arthrodesis in goalS produced similar fusion rates
and biomechanical stifFness when compared with
autogenous bone (Pintar et aI., 1994). 11 is interesting to note that the goat holds its head creCl, and
thus loads the cervical spine similarly to human
bipeds. Goat models for cervical spinebiologicallbiomechanical testing are, therefore. populal:
A can inc thoracic anterior al1hrodcsis model
yielded contrary results when tested biomcchanically:
autogenolls iliac crest grafts were stiffer in all motions
than \\lcre ceramic graft substitutes (Emery et aI..
1996). aturally grown coral is also a useful bone graft
substitute once it has been processed. Jt is available
con1rncl'cially in t\\-'o different porosities (200 ~m and
500 lJ.m pore size). The grafls of lower porosity had a
compl"Cssive strength comparable \vith biconical iliac
crest grafts, although they were much more briule.
They can therefore be considered appropriate for clinical lise with rcspect to their compressive strength
(Table 11-1). An increasing variety of prosthctic intcrbody cages ~\re becoming available (Shono ct aL, 1993);
long-tellll clinical results arc eagerly awaited. Additional anteJior (Fig. 11 ~31) or posterior intcI11al ft:-.:atio!l
using plates and screws for added support ma.\' be lIsed.
Posterior arthrodeses of the ccndcal spine
commonl:v performed after trauma but may also
L1sed to treat degenerative, inflammatory, or n
plastic conditions. Unlike anterior fixation devi
which arc solel y lIsed in the subaxial cervical sp
posterior fixation devices can extend up to the
ciput (Fig. 11 ~32). Posterior fi.'\ntion has commo
consisted of various wiring methods, which
used as a tcnsion band posteriorly (Sutterlin et
1988). Within the past decade. however, techniq
utilizing screws in both thc atlantoaxial and sub
ial spine have become rnorc popular.
Techniques for aLiantoaxial fixation have evol
from onlay bone grafting alone to cver~more sop
ticated methods of internal fixalion. Interlam
clamp fixation is clinically reliable and
(iVIoskovich & Crockard, 1992). Screw fixation te
niques may confer increased translational and a
rotational stability, but with greater surgical ri
A biomechanical evaluation of four commonly u
techniques for posterior atlantoaxial fixation
performed (Grob et aI., 1992\ Cadaver CI-C2 f,
tional spinal units were tested in nexion~cxtensi
lateral bending, and rotation, while intact as we
after a complete ligamel1lous injul)/. The fixa
techniques used wcrc:
l
I. Sublaminar wire with one median graft
(Gallie type)
2. \Vire fixation with two bilateral grafts
(Brooks lype)
3. Transarticular screw fixation (Mager!)
4. Two bilateral posterior Halifax clamps (Fig
11-33)
str<:ltcd similar 01' irnproved fusion rates com
with stand-alolle bone grafts.
Anolher group <:lssl."s;ed lilt, biollH.:L·hallical s
it~· 01" sen.'n different cervical n:construction
ods using 24 c(lif cClyieal spinl' segments (KotC
al.. 1994). Thl'~' studied Ihn:\:.' posk'rior-nnl,\'
niques. anlcrior-onl.\' u...· lyical platL·s. antcrior
graft alone. and a combined <:lIlterior plate wilh
(erior triple wiring. The results dcmollstnlled
for onL'-lc\"L'1 instrlbilit\', the stiffncss \"alues
consllliciion with an a;1terior procedure alOlle
si!.!nificanth- smaller than those of all the othL'
st;ucts tcst~'d. Intcrl'slin!!lv, stiffness w<:\s n.:sto
prc-in,iul-y le\'cb in specin'1L'nS \\'itlt lwo-lc\'L'! i
patterns reconstrucled using an~' of the three
riot" 1l'lL,thocls. TilL' findings do not SlIpport the
C':\clusi\'L' ank'rior methods in either posteri
An example of an anterior cervical plate applied over two
motion segments on a model of the cervical spine (Peak
cervical plate, DePuy Acromed, Inc., Raynham, MA). The
screws do not penetrate the posterior cortex and are
locked to the plate to prevent backing out.
The Gallie technique allowed signincanLl.y morc
rotation in Oexion, extension, axial rotation, and lateral bending than did the other three fixation techniques, No significant differences were noted in the
amount of rotation between the other three fixation
techniques, although the Magerl transarticular
screw fixation tended to permit the least amount 01'
rotation, as might be expected,
A variety of anterior cervical plating systems are
currently available, and many have been subject
to biomechanical analyses. Investigators demonstrated that in a single-le\'e1 procedure, an anterior
cel"Vical plate serves as a load-sharing device rather
than as a load-shielding device, enabling graft consolidation as observed in other clinical studies
(Rapoff el aI., 1999). Bone lInion may be expecled to
occur at a lower rate if the bones are shielded from
compressive forces; however, clinical experience
with anterior cervical plates has generally demon-r
EIIEI
Posterior occipitocervical fixation using a highly mod
stainless steel loop (Ransford loop, Surgicraft; Redditc
Worcs" U.K.) wired into place on a skeleton model. T
cranial loop is attilched to the occiput Llsing wires tha
pass through paired full,thickness burr holes. The wir
not passed around the foramen magnum. The loop is
bent to conform to the posterior craniocervical angle
may be adjtlsted intraoperatively. The limbs of the lo
are attached via sublaminar wires at each level. Note
the axis laminar wires are tightened below the flare
loop, which serves to maintain distraction between th
ciput and the axis vertebra. Sublaminar wire use is on
many techniqlles available to achieve segmental post
fixiltion. Reprin:ec! (''1iIh permission from tvioskovirh R.. e
(2000i. OcopitorervJCaJ stabi/izallon for mj'eJopathj' In pat
: ','Itil rlleumdtoid (1f!hritI5 ImpliratlotlS 0; /lO! ball£' gr(lf;lflg
BOllc JOlni Sur£). a2A. 349-365
!
•
C1
C2
Diagram of a skeletonized atlantoaxial motion segment
viewed from its posterior aspect. The method of applica~
tion of the interlaminar clamps can be seen. The bone
grafts lie under the Halifax clamps and allow the posterior
construct to be locked tightly, providing immediate stability. Reprinted with permission from rvloskovich, R., & (rockard,
HA (992). Atlantoaxial anhrodesis using inrcr!aminc1{ clamps.
An improved technique. Spine, 17, 261-267
threc-colurnn instabilit.y. In three-column or multilevel instability, combined front and back procedures utilizing an anterior plate and posterior triple
wiring demonstrated clear biomechanical advantages. Clinical use of anterior plate fixation without
posterior fixation for three-column cen/ical injuries,
however, resulted in satisfactory clinical outcomes
(Ripa et aI., 1991). These results underscore the
need for good clinical evaluation and studies that
take into account the fourth-dimension-time. Fu
sion is a biological process that occurs over time
and supercedes the importance of in vitro or
computer-simulated biomechanical studies.
w
equipped vehicle (King & Yang, 1995). Flowever,
long after airbag devices becan'lC available, air
injuries involving front seat passengers began to
described and by late 1997, 49 child deaths and
seriolls injuries had been attributed to passeng
side airbags (Marshall et aI., 1998).
A retrospective review was performed by three
diatric radiologists to determine if a pattern of inj
was common to this new mechanism of pedia
trauma (Marshall et aI., 1998). Their findings, as w
as the findings of other studies, demonstra
that passenger-side airbags pose a lethal threat
children riding in the front seat of an automo
(Giguere et aI., 1998; McCaffrey et aI., 19
Mohamed & Banerjee, 1998). Many of the child
killed in these accidents \vere front seat passeng
involved in low-speed collisions in \vhich the dr
sustained no or only minor injuries; the pattern
injury' seen is different in the rear-facing infant
seat versus the fOl\vard-facing child car seat. In
formec the injury' to the infant \Vas often mas
skull injury and cerebral hernatoma as a result of
proxirnit)' of their heads to the airbag, while in the
ter, children sustained man)' more cervical injur
'1\\'0 of the older children had autopsy findings o
lanto-occipital dislocation and one sllstained a "n
decapitation" injury, demonstrating the vulnerab
of the pediatric cervical spine to the explosive fo
of an expanding airbag that hyperextends the ch
fragile neck. Recent guidelines from the NHTSA (
11-3) were prompted by reports of airbag injurie
children and emphasize that children of an}'
should be properly secured in the back seat (Natio
Highway Traffic Safety Administration, 1996).
A mathematical simulation was performed
study the potential of head and neck injury to an
belted driver restrained by an airbag (Yang et
Nai;ional HighwayTraffic
Safety Administration
.Guidelines Regarding
Airbags & Children
BIOMECHANICS OF CERVICAL TRAUMA
Airbag Injuries
Motor vehicle accidents continue to be the leading
cause of injury-related deaths in the United States.
[n 1984, the National Highway Traffic Safety Adrninistration (NHTSA) required that automatic occupant protection devices (airbags or automatic
scat belts) be placed in all automobiles in the
1987-1990 model years. In 1993, the passengelcside
airbag was introduced. Studies generally' concluded
that front seat occupants are adequately! protected
against frontal impact if belts are worn in an airbag-
The back seat is the safest place for children of any age
to ride.
Never put an infant (less than 1 year old) in the front of
a c"ar with a passenger-side airbag.
Infants must always ride in the back seat, facing the
rear of the car.
Make sure everyone is buckled up. Unbuckled occupants can be hurt or killed by an airbag.
1992). It was found lhal when the standard 20° angie steering wheel was used, neck joint torques were
decreased by 22%. The resultant head acceleration
increased 41% fTom the baseline study when a vertical steering \vheel \\"las used. "If the verlical (limensian of the airbag was reduced by 100/0, neck joint
torques were increased by l4%, while head acceleration showed a slight decrease of 90/0. Although
ideal dimensions and inflation rates for airbags remain elusive, their use has resulted in a significant
reduction in head and neck injuries.
Whiplash Syndrome
Whiplash syndrome is a complex se' of symp,oms ,hal
may present after an acceleration hyperextension injury. These injllIies typically occur when a Car is struck
From behind, bUl may also be caused by lalcml or
Fron'al collisions (Barnslev el aI., 1994) (Case Study
11-2). The acceleration of the car seat pushcs the torso
of the occupant forward with the result that the unsupp0l1ed head falls backward. resulting in an extension strain to the neck. A secondaJ)' Aexion injlll)' may
occur if the vehicle just struck then strikes another vehicle in front and just as suddenly decelerates again,
throwing the occupant forward once more. Crowe
coined the tellll 'whiplash' in 1928 in a lecture on neck
injllIies caused by rear-end automobile collisions in the
United States but latcr reported that he regretted using
the term (Breck & Van Norman, 1971) because il describes only the manner in which a head was moved
suddenly to produce a sprain in the neck and not a specific injUl}' pattern. Therefore, we refer to this clinical
entity as whiplash syndrome, not as whiplash injllry.
Allhough whiplash injuries are a common traumatic event, the pathology is poorlv understood. 01'len the severity of the whiplash trauma does not
con-elate with the seriousness or the dinical problem, which can include neck and shoulder pain,
dizziness, headache, and blurring of vision (Brault
et al., 1998; Ettlin Cl aI., 1992; Panjabi el aI., 1998a;
Sturzenegger el 'II., 1994). Apart Ii'om a frequently
observed loss of physiological lordosis, a radiographic examination of the cervical spine is often
normal. Even newer technology such as MRI is not
always able (0 I-eveal a soft tissue injul}'. MRI examination of the cerebrum and cervical spinal column
perrormed 2 days after a \vhiplash neck sprain inJUI)' in 40 patients did not deli~cl an)' pathology con~
nected lO the injlll)', nor was the MRI able
predicl
symptom development or outcome (Borchgre\fink d
aI., 1997). Injuries thaI have been documented in-
'0
elude inlerspinous ligamcllt tears, spinous pr
fractures. disc rupture. ligamcntum f1avulll nlp
facd joinl disruptiun, and stretching of 11K.' ~\I
musdl's. The diagnosis and managcll1clH of
lash injuries an.. . oflel1 t.:ol1l"oUIl(!L·d b~r concom
psychosoci;':il ,lilt! IlH.'dicolcgal isSlIl..."S':IS well (\
et ,Ii., 1998j.
One of the earlv (unpublishcd) biomceha
studies or whiplash injury was clone b.v the la
Irving Tuell. nil orthopuedic surgeon in Se
\V~\shillgton. I-h.: uscd a cinl' camenl to photog
himsdf driving as he was rammed from behin
his surgical J'csidL'nt driving anothLT car. Fr
drawn frolll Lhat movit.' ckarly demonstrate til
pen:xtension of his neck over the sl.'alback of h
(Fig. 11-34). One C;:\l\ also scc thl' dlect of in
l"orces on tht.: mandible as his jaw snaps opt'n
the Hcccicratioll forces his !lead backwards.
mechanism Illi.l\ explain the tL'mporomandi
injuries thal (In,.' a COllllllon accompanilllL'nt o
\'ical \\'hipbsh injuries_
r\ recent sLud~' of rcar·end collisions qual
!i\'d~' elucidated the m.:tllal neck movements
wke place (Castro CI aI., 1997). A1kr
ramllled from behind, the Illcan accl'ler<:\tion
target \'chicles was from 2.\ to 3_6 g. Maxim
tension was n.'ached when the head contacte
headrest; the angle b('!wt...·L·n the head and t
bOth' varied From 10 to 47" (meall = 20"). In th
SCIlCL' or a headrest, tht..., maximal r('cordcd e
sion was 80". FollOW-LIp clinical and rvIRI cxam
tions were pcrformt.:d_ The sttlcl~! concluded
the "limit or hannlessllcss" for stresses ar
from rear-end impacts with regard to the ve
changes lies bd\\'cen 10 and 15 km/h, In stud
reproducible \\-hiplash trauma mode! using w
cervical spinc specimens 1ll00Jnh.:d on a belH.:
sled was used to simulate rear-end collision
increasing horizontal accderations applied t
sled (Panjabi ct 'II" 1998a.b). Bo,h sled and
kinematics CHIl be measur'cd using potentiom
and acceleroll1l.~tcrs. Using this whiplash modt
S-shapL'd cun:c was (kscrihed in whiplash in
in which the lowt.:'r cervical spinC' h.\'perext
~JIld the upper cCITical spine flexed (Grauer
19(7). The invesLigators r~h that Ihe in.iur~' w
currcd during the hyperextension phase i
knvcr cervical spinc,
Correct positioning or adjustable headn..'s
hind the skull. not hehind the Ileck, is impo
Tests using a H~'gc sk~d and a Hybrid III dll
wcre pcrrorl11L:d by other iJ1\'estigalOrs to deter
fi
.•,;;;;: _ - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - ------------ ------------------------.-
,
;"
;:
.. ,
'
,~
if ! Whiplash Syndrome
;::',:'i::'/.::,';):
::'»::)/:,:'':-''P,}\
32-year·old man was injured in a car accident. The trau-
::~9>:~;:_~i~;~ matic event happened after being struck from behind. [n
/~~{~:d{t:ih.iS
case. the inertia of the head and the flexible spine re-
from Fr(1nkel, VH. (1972). Whiplash injuries to the neck. In
C. Hirsh & Y Zotterman (Eds.), Cervical Pain (pp. 97~111).
New York: Pergamon Press.
"?g;'sulted in marked hyperextension movement of the head. The
.¥a~celeration of the car seat pushed the torso of the occupant
:,~?;~
;7'''':;
$Jo~va{d with the result that the unsupported head
'ft~~ard, resulting in an extension strain to the neck.
fell back-
lr -;.
The palient presented with severe neck and shoulder pain
accompanied by sleep disturbance. The right sternocleido-
j-
mastoid muscle was approximately twice as large as the left
sternocleidomastoid muscle. After muscle testing against re-
t,
>.:
..t.
' f..
~>
sjstance, the pain increased (Fig. cs 11-2-1).
'To decelerate the posterior rotating head, a moment and
g>:a.force must be developed by active and passive stabilizers,
i.-?int surfaces,
and the intervertebral disc. The moment and
force give rise to tension, compression, and shear stresses
.',: .find strains in various parts oi the neck Cdusing, damage.
fl., During the collision, rhe muscle tension increases with ihe
, velocity of lengthening. A tension inappropriate for the
length tension curve occurs and partial rupture of the sterno4
cleidomastoid muscle is produced. Reprinted with permission
Case Study Figure 11-2-1.
'.- -..
I
I
I
I'IDIIDI
1~_H.a_nd_-_d_r.aw_n.c.e.I .,
•
~.
,;
.. ...
,.-
B
(A) Before and (B).tr.a.c.ed_f'.o.m_t_w_o_f.r_am_e_'_o_f_a.C.in.e.-.m_o_V_ie_o.f.D.r_.I.rV.i_n9_T_u_e_lI_d.,.iV.in.9_h.i'.c.a.raher bein9 rammed from behind by another car driven by hi, re,ident_
_
biomechanical responses for the various conditions
observed in normal driving (Viano & Gargan. 1996).
Risk of injury was assumed to be proportional to
neck extension. A low headrest position carries a
relative injl1r~' risk of 3.4 in rear-end crashes. compared with 1.0 for the favorable condition. If all adjustable headrests were placed in the up position,
the relative risk would be lowered to 2.4, a 28.30/0 reduction in whiplash injury risk.
Significantly more complex acceleration injuries
ascribed to high G-force activities are now occurring.
There is a report in the Htel·ature of cervical injury in
pilots of F- t 6 fighter planes, citing a I-year prevalence of neck injury of 56.6% and a career prevalence
of neck injury of 85.4% (Albano & Stanford. 1998).
Maintenance of neurological homeostasis and
protection of the spinal cord. nerves. and vessels
and support and protection of the skull are the ultimate tasks of the cervical spine. An appreciation of
the principles presented should afford a greater understanding to physicians and allied health professionals involved in the treatment of cervical spine
pathology. 'vVith increasing technological advances
in our society, \ve remain vulnerable not only to
common types of cervical trauma but also to idiosyncratic methods of injury' related to these new
technologies. Acceleration injuries have been OCCUI'l'ing for decades but cervical spine injuries caused
by airbags have only recently been described, and
aliI' involvement in increasingly greater speeds and
higher risk activities places LIS in increasing jeopardy, As our body' of knowledge expands, we continue to discover new methods of injury. It is critical
to pursue rational treatments for these disorders
based on sound biomechanical principles,
5 11111111 a Iy
1 A functional spinal unit or motion segment
consists of t\\'o adjacent vcnebrae and the intervening intervertebral discs and ligaments between the
vertebrae.
_2 Whcn eithcr rotation or translation of a body
along one axis is consistently associated with a simultaneous rotation or translation along another
axis. the motions arc coupled.
3,; Intervertebral disks cxhibit viscoelastic properties (creep and relaxation) and hysteresis.
(4;' Discs arc able to withstand greater than nor-
l1l~r loads when compressive forces arc rapidly ap-
plied, which protects the disc from catastrophic
lIrL' ulltil e . . . trl"lllcl~· high loads are applied.
5 Vi".'I·lL"bl·al bod~'
from upper ct.'n:ical
cUlllpr~ssioll str~ngth
10
incre
IO\\'L'r lumbar lev-cis.
6 TIlL' mean sagillal diametL'r of the malt.- a
spinal canal for C3-C7 is close LO 14 Illm; the sp
cord diameter is about 10 111 Ill.
7 LigamclltuJ11 flavlIJ11 is undt..'1' tension
when the spine is in a Ill:1Itral position or somew
cxtended. prestressing the disc and pro\'iding s
intrinsic support 10 tht..' spine.
S iVlllscles pla~' a critical role in b~lsic post
homeostasis. Patients with paral~'zL'd ct..'lyical m
cles arc unable to support thdr heads.
9 The spinal cord has some longilUdinal c1a
it.'· blll it tole-ratL's axial translation pond.". It is
translatolY forces {hat t.\'picall,\· rcsult in neurol
cal injul·.\',
10 Instant c('ntl.'!' anal.\'sis indicates that tange
lnotion (gliding) wkcs placl.· bel\\,(.'c·]1 lltt::.' facet jo
as lItc cervical spine is flexed alld extended. The
or the intelTcrtebr~il !"()ramina inc.TeDscs with fle
and deCl'l:ascs \\'ith extension,
11 Kinematic instability' refl'rs to the quantit
motion (too much or too little) or till.' quality of
tion present (alterations in lile normal pattern
both. Component instahilil.\' addrcsses the clin
biomechanic~ll role of till.' \'ariolls anatomic st
tures of the functional spinal unit.
12 An." motion s~~glllent in which all of the a
rior ek'llkllts or all the posterior ekments m·e
stn>yed or are unabk' to function should be con
ered unstable.
13 A significant incn:asc in annulus stresses
scgrnclltal mobility' may' OCCUI' when bilateral f
resection exceeds 50(*;.
14 Appropriate use or internal fi:G1lion helps L
(:reasc the fusion rate and maintain structural a
ment.
15 Front seat occupants are adcqLH.llcl~' prote
frontal impact if scat belts arc worn in
airbag equipped ,·chick. Passenger-side air
pose a lethal threat to children riding in the f
scat 01" an automobile.
~lgain.s{
16 \'Vhiplash s.\"ndrome is a complex set of s~
tOI11S thal may present aher an (\("ccleration hy
extension injury.
REFERENCES
Adams .•\LA. 8.: Hutton, W.e. (1980). The ('freet 01" poslllrl' on
the rok of the apophysial joints in resisting in[e!'n:"rlebral
compressive forces. J BOIlt' .Ioill! SUfg. 62B, 358-62.
Albano, 1.1. &. Stanford, J.B. (1998). Pn.'venlioll of minor neck
injuries in F-l6 pilots. ill·juf Space EI/I'iro!l ,lied, 69.
1193-·1199.
Bailcy. R.\V. (1963). ObSCIT<ltions of cerdcal inten"crll'bral
disc lesions in fraclUres and dislocations. } BOlle Joillf
Sutg" ':'/5..\, 46l.
Barnsley, L, Lord, S" &. Bogduk, N. (1994). \Vhiplash injun',
Pail/, 58. 283-307.
.
Bell, G.I-!. (1967). Variation in tht stl't.'n~[h of vcrtdH"<! with
;lge and their relation to oSkoporosi:. C(/Ici( Tissl/i.' Res,
/, 75.
Bennett, S.E., Schenk, RJ., & Simmons, E.D. (2000). Active
range of motion utilized in the cervical spine to perform
daily functional tasks (Abstract). Associalioll
BOlle alld
Joillt Sllri~colls 5llld .'lIlII/W! J1,-,<-'tillg. San Antonio. TX.
\lay 2000.
Blackley, H.R .. Plank, L.D .. & Robertson, P.A. (1999). Determining the sagillal dimensions of the canal of the cervical
spine. The reliability of ratios of an~ltomical Tlleasure-
or
ments.JBollcJ()intSlIrg,8IB.110~112.
Borchgrevink. G.. Smevik. 0 .. Haave. I.. et al. (1997). \lRI of
cerebrum and cer\'ic<:ll cOIUTlln;:l within two davs after
whiplash neck sprain injury. Ilijury. 28, 331-335 ..
Brault. J.R .. Wheeler. J.B., Siegmund. G.P.. et al. (1998). Clinical response of human subjects to rear-end automobile
collisions [published erratum appears in Arch Ph"s :\led
Rehabil 1998 Jun;79(6):723j. Arch Phi'S "II:'(/ RellO'hi! 79
72-80.
. ' ,
Breck. L.\V. & Van Norman, R.W. (1971). \lcdicolegal aspects
of cervical spine sprains. C!ill OrtIJop, 7.J, 124~8.
Breig. A.. Turnbcll, I., & Hassler. O. (1966). Effect of mechanical stress on the spinal cord in cen'ical spondylosis.
J NCllwsllrg, 25. 45-56.
.
Cart:!', n.R. & Hayes. \V.c. (1977). The compressi\'(.' behm'ior
o! bone as a two-phase porous structure. J BOIIC Joint
SlIrg, 59:1. 954-962.
Castro. \V.H .. Schilgen. :\1., :\Ieyer, S., et a1. (1997). Do
"whiplash in,iuries" occur in low-speed rear impacts? EliI'
Spill" J, 6, 366-375.
Cherubino. P.. Benazzo. F., Borromco, U., et a1. (1990), Degenerati\'e arthritis of the adjacent spinal ioint~ follo\~'irH!
anterior cen'ical spinal fusion: Clinicoracfiolouic and sta~
tislical correlations. Iial J Orthop 7/"a/l/lUlIO!. 16. 533-543.
Clausell, J.D .. Gocl, \'.K .. Traynclis. V.C., el a!. (1997). Uncin;:tte processes and Luschka joints influence the biomechanics of the ccrvical spine: Quantification using a finite
dement model of the C5-C6 segment. J Orrhop ~Rl's. /5.
342-347.
Cusick. J.E. Pintar, EA .. & 't'oganandan. N. (1995). Biomechanical alterations induced bv Illultilevel cervical
laminectomy. Spille, 20. 2392-2398:
Denis, F. (1983). The three column spine and its significance
in the classification of acute thoracolumbar ~pinal injuries. Spille, 8. 817-831.
Dickman, C.A .. Zabramski. L\1.. Hadley. ,\1.N., et al. (1991).
Pediatric spirwl cord injury without radiographic abnor-'
malities: Report of 26 cases and review of the literature . .I
Spillal Disonl. .J. 296-305.
Dvorak, J .. Froehlich. 0 .. Pt.:llning. L., et a!. (1988b). Fun
~ional r:ldiographic diagnosis of the cervical spine: Fle
Ion/extension. Spillt'. 13(7).748-755.
DnJnlk, .I .. f·la~'ek. J., &. Zehnder. R. (1987). CT-functional d
agnostics of the rotatory instability of the upper cervic
spine. Part II. An evaluation of healthy adults and patien
with suspected insl;:lbilit.v. Spillt', 11(8). 726~731.
Dn,lt';lk, J. &. Panjabi. \1.\-1 .. (1987). Functional anatomy
the alar ligaments. Spille, /2, 183-189.
'
Dvorak, J., Schneider. E.. Saldinger. P.. et al. (1988a). Biom
chanics of the craniocervical region: The alar and tran
verse ligaments. J Orthop Res. 6,~452-461.
Emery'. S.E .. Fuller, D.A .. & Stevenson. S. (1996). Ceramic a
{eri(~r spinal fusion. Biologic and biolTlcdwnic<l1 compar
son In a canine Illodel. Spilil'. 2/. 2713-2719.
Ettlin. T.iVI.. Kischka. U.. Reichmann, S., et al. (1992). Cer
bral Sy'lllptoms after \\'hiplash injun' of the neck:
prospective clinical and neuropsycl{ological studv
whiplash injury. J ""'t'/lro! ,Vellrosllrg Ps\'c!Jiottv.· 5
943-948.
'
.
.
FeipeL V.. Rondelet. B., Le Palkc. J., et al. (1999). Norm
global motion of the cervical spine: An declt'O!!oniometr
study. Ctill Bill!IICch, /.J, 462-470.
Fidding. J.\V. (1957). Cinero<.:ntgenograph~· of the norm
cen'ical spine. J BOIlt' Joilll SlIr~. 39:\. 1280-1288.
Fielding. J.W.. Cochrnn, G.\'.B .. La~~'sing, J.F. !II, et a!. (1974
Tears of the tranS\'erse ligamenl of the atlas: A clinical an
biomechanical study'. J BOlle Joil/t Slitg. 56:\. 1683-1691
Fid~ling,. J.\V.. Hawkins. R.J .. & Ratzan. S.:\. (1976). Spine f
sIon lor atlanto-axial instability. J BOIIl' Joilll Slirg. 58
400-407.
.
Fountain. F.P.. :\linear. \V.L., & Allison. R.D. (1966J. Functio
of [on gus colli and longissimus cen'icis muscles in ma
Arch Phys Alcd. 47. 665'::-669.
.
Frankel. \'.H. (1972). \Vhiplash injuries to the neck. (n
Hirsh 8.:: Y. ZOlLerman (Eds.). Ct'n'ica{ Paill (pp. 97-111
New 'I'ork: Per£amon Press.
Fuller. D.A .. Kirk()alrick. J.S .. Emery, S.E., et al. (1998).
kinematic study of the cervical spine before and after se
mental arthrodesis. SpillC, 23, 1649-1656.
Giguere, J.E. St-Vil. D.. TUrlnel. A.. et 'II. (1998). Airbags an
children: A spectrum of C-spine injuries. J Pcdioll': SlIn
33,811-816.
Gosch, H.H .. Gooding. E... & Schneider. R.C. (1972). An e
perimental study of cerdcal spine and cord injuries.
trill/nUl. 12.570-575.
Grauer. J.N .. Panjabi, \1.M., Cholewicki. J .. et al. (1997
\Vhiplash produces <In S-shaped cun',Hure of the neck wi
hyperextension at [ower le\'cls. Spilw. 22. 2489-2494.
Grob, D., Crisco, J.J. 11(, Panjabi, \1.\-1., et al. (1992). Blom
chanical nalu,ltion of four different posterior atlantoaxi
fixation techniques. Spillt', 17(5). 480-490.
Hansson. T. 8.:: Roos. B. (1986), A!!c chan!!es in thc bone mi
eral of the lumbar spine in n~}rnwl \~Olllen. Calcit' Tiss
llIl, 38, 249-251.
I-lansson, T., Roos, B., & Nachemsoll. A. (1980). The bo
mineral .content and ultimate compressive strength
lumbar vertebrae. Spint', 5. 46--55.
~
Harms-Ringdahl. K. (1986), On assessment of shoulder eXl
cise and load-dicited pain in the cervical spine. Bio}//
dUlI/ica! AIU/!ysis oj' Load-ElfG-,I/t'thodoiogica! St/U!i
u(Poill Pnl\'oked by Ex/reille POSilioll. Karolin~ka Inslitut
Universit~' of Stockholm.
Herkowitz. H.N. ( 1988). .-\ comparison of ~ln".:rinr cer\'ic~ll fu·
sion, (:cr\'ic~t1 bmilll..'ctollly, and cc:rviC<11 hllninopl:lsty for
the surgical managi:nH:nt of Illultipli: kvd spondyloti\'"
radicul()path~'. 5pillt',
13, 774-780.
Herzog, R.J .. Witns. J.J .. Dillingham. M.F.,l'l al. (1991). NorJ1l~d (:elyiGII spinc morphomelry iwd ce.:rviL"ul spin:lI slcno·
si:-: in ,lsymplOmatic professional foot hall players. Phlin
film r~diography, muhiplanar computed lomogr''11hy, and
m,lgn\,.,tk n:sonancc imaging. 5pil/('. 16, S178-S186,
Hilibrand, A,S .. Carlson, G.D" Palumbo, M.A .. t:t al. (1999),
Rndiculop.l\hy and myelopathy at sl.'gmcnts ;loj.u;l:l1t to
lhc SiL\.' of'l previolls anterio!" cc:n'ic~tl anhrodc:.sis, 1 BOllt'
10illl Sllr~, 81A, 519-2:>,
Holdswor(h,
(1963). Fr<lctures, dislocations, :lud
lurc-dislocOitions of thc spinc. 1 BOlle 1uiut SlIrg,
r.w.
6-10.
Huntl.'r, L., Braunstein, E., & B~li)('y. R. (1980). Radiogr:'lphil;
chilng('s following nnterior fusion. Spille. 5, 399-401.
hrcho. ~'1. (1981). Calcium phosphale CI.:ramics as h~Il'd tissul.'
prosthctics. CJiIl Ort/lOp, 157,259-278.
Jenkins, O.H. (1973). Extensiv(' ccrvic~1 larnincclOlll~" LongIeI'm results. Bt'l S/fl"g, 60. 352-$54,
h:.~lz.lrl'lll. I...E. (1975}. Creep charactaistics of thl.' human
spinal t:oluIII11. Onhol' C!ill North Alii. 6, 3.
K"tl.:trinn, L. (19i2). Dynamic response Chan\l;leristi(,;s of Ihl.·
hUIlWll inlervt:rlebral column: An cxperirllc:nlal stud,\' of
autopsy specimens. Acta Orr/IO{J Sea lui, No, 146,
King, A,I. &: Y.Hlg, K.H. (1995), Research in biolncchanic:-I of
occupant pn.Hcclion, J hal/IIUI, 33. 570-576.
K.il:dl~lra, Y.. Uda, B .• & Tachihana. S. (1995), Effect of spin:d
cord stn:Lching uuc to head fk.\ioll on intr~lnH:duliary
pn:ssul'c. NI.!/lrol ,VIed Chi,., 35. 285-288.
Kot;'llli. Y., Cllllningh~un, B.W" Ahumi, K., el a1. (1994). l3io·
mechani<.:.d anaiysis of cervical st:ibiliz<\lioll systems, :\11
i1SSCSSll1Cllt of tl',l(lspedicllhlr screw fixation in the l.,'\":r\'ic;~\l
spinc. SpiJlt·. /9, 2529-2539.
Kriss, V.M, I.~ Kriss, T.e. (1996). SCIWORA (Spill~ll cord injury
without rlldiogr;'lphil; <lhnormality) in infnnts and chil·
dren. CJiIl Pt,·dii/II", 35, 119-124.
KUlll:ln:san, S., Yo~an;'lIldan. N.. PimaI', EA., et al. (199;). Finih..' ,·lell11.'!H modeling of cervical Iamillt:ctonw with
gr'l(k'd f.ICCtcClOI1lY. 1 S~)ill(J! Dison!, 10, -10--16.
.
Lonstcin, J.E. (1977). Posl.laminc:clomy k.... phosis. CliI, Or·
t/lOp, /28,93-100_
Lysdl. E. (1969). ,\-Iotion in thc cer\'ic~\1 spine. An eXIH.'rilllcn·
1<11 sludy on autopsy specimcns . .'lcla OnllOl' S,:cwd, SlIpp!
/23. 1-61.
Marshall. K.II'.. Koch. B.L.. & Egelhoff. J.e. (1998). Air bag.
rcl<\lcd dC<.lths and seriolls injuries in childrl.'n: Injury p;,H·
terns ~Ind imaging findings. A'/I J Neuwnlt!;o!, 19,
1599-607.
j\·kBroom, R.J., Ha\'t's. W.c., Edwards, W.T.. d nl. (1985).
Prl.'diction of \·c'-·Il.'bral body cOl1lprcs~in' fn.lClurc using
quantitath'c compuled tomogr'lphy. 1 BOlle loil1f Surf,:,
67A. 1106-1114.
McC.11'frcy, M., Gl.'rman, A., Lalondt:, F., (.'t al. (1999). Air bags
~IIH.I children: A pOlcntinlly I!.'lhal (ombin:ltioll. 1 I'l·dillir
On/lOp, /9, 60-64.
~'lil1lllnl,
M., tvtoriya, I-I.. Watanabe, T., ct al. (1989). Threl.'·
dimension,11 motion anal~-'sis of lhe cer\'ic:d Spillt: with
spl:cial n.:fcrcnct: to lhe axial rotation. Spille, 14,
113;-1139.
.\li.\:uaki, K.. T'I(b. h .. \lahIHb. Y.. l'[ :11. / 19ii9J. I'o:l·\ll·n:>i\'\.· silJlllll:llIl'(JIl~ lllulti:>l'gllll'lll:tl 11I:-lllll fC'I- l'l'!T
1JI.\dl'P~lthy widl l-l'l"\'il':c! ill~l:JhiliIY :tlltl k~phlJlil' :t1
S·:-hapnl ddorlilitil·~. Spi!ll'. 1,1{ 11 L 1160-1170.
\I<oh:Il11nl. :\.:\. & Ball\.·rin·. ,\. (199$1. P;lltl:l"lI:> of injun
:>oL"i;'lll'd wilh :llltullH,hik ;'lirh;lg Il~l·. J1o.'1.~nl/l .\I•.'d
4:15--i38.
.\({lrolll·\. S.P. &. Sl,.·llUh/, :\.B. (19$;). :\n:d.\:-i:- ~llIti Illl',h
flIl'llt (,I loads 1111 !Ill' nl·d. hew,' Onh R,'., Soc, J(J, _,1
.\l(Jn,nl'\, S.P., S.:hllllz...\"B .. ,\1ilb', 1.. l't :11, (I%SJ. l.t1:HI
placl'llll'lll pnJjh'rtil'~ 01 10\\'\.'1' n'l'\·;l.'al ~pirh' motion
ml'IlI~. J HiuIJlt'ch, 21, 769-779.
,\llI~k(l\'ich. R. (1 'J94J. ,\lhllllCI-;t\i:!l in~l:lhili[.\·, Spill,': SflI
fhi' .-\1"/ Uni<'l\'.,. S(31. 531-5-i9.
.\Io:-kmich, R. (1\)97). Cenk:d ill:>I'lhilit.\ (r!h'HIIl:
d\\:lrfi:-IIl, (kt!l'lJl'rati\l·. Illhl:[':-). In 1'..11. Brid\\\.·11 l\:
Dl·\bld (Ed:-. .i. Thl' 7~·.r1hfJc,J~ o( Spinll/ S/ft'J,:CIY
969-1009). Phil:tddphi~l: l.ippillCl'tt·Ran·1l Ptlhlislh-"I'~
.\to:>kll\-ich R.. 1,,'1 a!. {2000i. On'ipit(ICl'ITicti :>1:lhili/.'lIi(Jl
Illydllpalhy ill Jlatielll~ willi rhl'ulll;'l[oid :tnhrili:>. Im
ClliollS (If 1101 hOll\.~ gr:dtil\~. J Built· luhlf Sun.:,
1
349-365.
,\I(I~ko\"idl,
R, &. Crod.anl. 11.,\. (199~). :\tlanloa
:Irlhmlksis tlsinl~ illll'rblllin:lI' l'lalllp~. :\11 ill1pnj\"l'd
Iliqlll'. Sllill,', f7, 261··1C\1.
.\loskO\'kh, R. ~\: ./Olll'S, [)".\, (19~~), CpPl:r l'l'n ical spin
s[I"t!Il)('nl:ilioll. Spill": SieUc' ('I Iii,' '\0 Un'i,'I\'s,
2.,3·_·2:'3.
\losko\il:h. R., ShOll, S". & Zil:\n~. Z,ll. (1006), DOl'S till'
\'k,i1 '::1I1,L! to hod\' r:llllJ 11I"l'dil'l ,.;pill:d Sll'llOSi<i Bill! l
.luinr Oi:>, S..:;. 01-71.
\l.\'l,.'rs. B.S. l\: \Vinkl']sll'ill, 13. . \ , (1995l. Epidl'll1i(ll()g~, l'
l'iC:lliu!l, 1l1l'l'hanis1l1, alld lokr~1I1(l' (II' hUIll:1I1 l'('r
spinl,.· injuril:s. Crir Ra Biulllni 1-:11,::', 23, 307·_·+09.
\:adll,.'IllSUIL A, l\: E\":lllS . .I.H.ll\)tJ$J, S(Jllll' Illl'l.'h:1I1ic:d p
l'r1il'S of till' third hunwll IlImh:l!' inll,r!:llllin:lr lig;,u
lligallh.'1l11l1lJ f!;'lnlIII J. .ll3iuJllcch. /, 101.
\::ltillnal lIigh\\"a~' Trani..: S:tfl't.\" Adlllinbll':llioll. (1996)
h:lg akn. Allil LJllc'I~!,: ,\It'd, 2S, 14~.
\owinski. G.P.. Vi:-ariu:-. II.. \olk, I..P.. l·t al. (1993).:\
lIIl"l-klilicll comp:lri:»un or Cl'l"\'k:t1 lalllinopbsly :lIal
\·il·:III:tlllin\.'i,·lomy with pn1grl'::.:-i\·l' bn·I\,·d(llll.\. Spine
1995-1004.
Os,,-·nbach. R.t\.. l\.;. .\kne"lt.'~, A.II. {1"JS9). Spinal l'(}rd in
WilllllUI I"i\diog:raphit:: ... hl1()rfll<.llit~' ill (hildrl'n. !'e
.\"('IIHhCi.
15. 168-174,
Pang, D. &: f'ollad, I.E (19S~)' Spill:t!l.."ord ill.i\lr~· withou
diogr:lphil' ,dmorlll'llil.\ in L"lIildrl'JI-Thl' SCJ\\·OR.,\
drollll·.l li"lllllllll, 19. 6~-I-(l6-1,'
P:lni"bi, ,\I ..\L Cholc\\"id,i, J.. i\ihu, K.. l't;,)1. (I99~h). S
lation (If \\"hipla:.;h Ir:tlllll~t using w!l(,k (l'rdr:,1 :.;pinl' S
illll·n:.;. Spillc', 23. 17-14.
Panj:lhi. ,\1.:\1.. Chn]\.'wil"ki. J.. \!ihll, K .. l'1 :t1. {199:-bi. Ca
br lig~1Il1l'1l1 strt'lchl's during ill \'Itro whiplash :o;i
li(IIl:-.l Spimd Dison/. 1/, 21i-l.'1.
Panjabi. \L\1., I\.r~lig. .\1.11 .. & God. V.I\.. (19~I). A 1l'.... hn
for m\.'i.t:'url·IlH.'llt alld dl'sl"l'iptioll of lhrl:('-diml·ll::.iolla
dl..'g:rl'I.'-of-frl'l'dolll motioll 01':1 body joint \d[h :1Il app
lion totltl' hum,1l1 ~Jlinl:. 1 WIlIJh'dltJlJics. 1-1. -i-l7--i60
P:llIjahi, M.M,. SUlllllll'r:-, 0.1.. Pdkt'r, R.R .. l't al. (19
Thr('l..'-d;l1ll'lISion:,lloitd-displ:'I.'l,.'l1ll'lll CIll"\"('S dui..' LO fo
(111 till' l,.'l'l"\·il'al spilll' . .f On/lOp Uf.", -I, IS~-161.
Paniabi, ,\1.,\-1., \Vhile, A.A. III. & Johnson, R.,\1. (1973). Cer\:ical spine mechanics ,IS a I"unction 01" transection of cornpOneJllS. 1 HiOlllt'c!/(lIIics, 8, 327-336.
Penning, L. & Wilmink, J.T. (1987). Rotation of the cCl'vical
spine. A study in normal people. Spille 12, 732-738.
Perry, J. & Nickel. V.L. (1939). Total cervical-spine fusion for
neck par,t1ysls. 1 BOIIt, loilll SIIJ'g, .Jl:\, 37-60.
Pintar, FA., \himan, D.J., Hollowell. J.P., L't 'II. (1994). Fusion rak and biorHechanical stillness of hydroxylapatite
versus autogenous bone grafts for antnior discectorny. An
in vivo animal study. Spille, 19, 2324-2328.
Pintar, F.A., Yoganandan, N., Pesigan, ~'1., L'l a!' (1995). Cervical vertebra! strain llleasuremellts under axial and eccentric loading. 1 Bioll/cch Ellg, /17, 474--478.
RapolT, A.J., O'Brien, T.J., Ghanayem, A.J., L'l al. (1999). Anterior cervical graft and plate load sharing. J Spillill Dis,
12,45-49.
Raynor, R.B., Moskovich, R., Zidel. P., et a!' (1987). Alteration
in primary and coupled neck motions after fncell.'ctomy.
NellnlslIl'.i.iery, 2l, 681-687.
Ri:ill. J. (1960). Effects of flexion-exti:lIsion movement of thi:
head and spine on the spinal cord and nerve roots . .I ,VCIIml ,VClII'OSlIr.i.i Psychi(/IJ', 23, 214-221.
Ripa, D.R., Kow,dl, \I.G., "'!eyer, P.R. Jr., et al. (1991). Series
of ninety-t\\·o traumatic cervical spine injuries stabilized
with anterior ASIF pl,lte fusion technique. Spille, 16(3),
46-55.
Rolander, S.D. (1966). \-lotion of the lumbar spine with special refi:rence to thi: stabilizing (.'ffect of posterior fusion.
An experimental study on autopsy specimcns. /lcw Orthop
Snllld,90, 1-144.
Roozmon, P., Gracovetsky, S.A., Gouw, G.J., ct a!' (1993). Examining motion in the cervical spini:. I: Imaging systems
and measurement ti:chniqui:s. 1 Bio/lled Ellg, 15,5-12.
Saito, '1'., 't'amamuro, '1'., Shikata, J., et a!' (1991). Analysis
and prevention of spinal column deformit.;.' following cervical laminectomy. I. Pathogenetic analysis of postlamini:ctomy deformitii:s. Spilll.', 16, 494-502.
Schneider, R.e., Chi:rry, G., & Pantek, H. (1954). The syndrome of acute ci:ntral cervical spinal cord injury with
special reference to the mechanism involved in hyperextension injuries of the cervical spinto J /":I.'ZlroslIJ'j;, 11,
546-577.
Schultz, A., Anderson, G., ()rtengri:n, R., et al. (1982). Loads
on the lumbar spine: Validation of a biol1ledwnical anal\'sis by measurements 01" intl'adiscal pressure and myoell'ctric signals. J BOllI.' .Ioillt SlIrg, 6",.1, 713-720.
Shono, Y., ivlcAI"ee, P.C., Cunningham, B.\V" et at. (1993), A
biomeclwnical analysis of decompression and reconstruction methods in the cervical spine. Emphasis on a carbonfiber-compositc cagt. } BOllI.' .loilZI S/lrg, 7504, 1674-1684.
Simmons, E.H. & Bradlev, D.D. (1988). Nturo-myorat
flexion deformities (/ the cervical spine. Spille,
756-762.
Steel. H.ll. (1968). Anatomical and mechanical conside
tions of atlanto-axial articulation. In Proceedings of
American Orthopaedic Association. J BOlle Joil/I SlIrg, 5
1481-1482.
Sturzenegger, .\1., DiStel"ano, G.. Radano\', B.P., et a!' (199
Presenting symptoms and signs after whiplash injury: T
influence of accident lllechanisms. NCllrology, .J.J, 688-6
Sutterlin, C.E., McAfee, P.e., \Vardell, K.E., ct al. (1988)
biomechanical evaluation 01" celYlcal spinal stabilizat
methods in a bovine model. Static and cyclical loadi
Spillt', 13, 793-802.
Takebe, K" Vitti, :v1., & Basrnajian, J.\!. (1974). The functi
of semispinalis capitis and spli:nitls capitis muscles:
electromyographic stud.v. Allat Rec, f79, 477-480.
Viano, D.e. &. Gargan, il,'l.F. (1996). Headrest position dur
normal driving: Implication to neck injury risk in r
crashes. Accidellt Allal fln'I', 28, 663-674.
Yoo. I..M .. Kumar"san. S .. Yoganandan. N.. 01 at. (1997).
nite elemi:nt anal~'sis of ci:rvical facctectomy. Spille,
964·-969.
Wallis, B.J., Lord, S.M., Barnsky, L., et al. (1998J. The p
chological profiles of patiellts with whiplash-associa
headache. Cephalalgia, IS, 101-105.
WertH.', S. (1937), Studies in spontaneous alIas dislocati
Acta Ort/lOp Sewld, 23 Suppl.
White, A.A. Ill, Johnson, R.~L Panjabi, :\1.M., et al. (197
Biornechanical analvsis of clinical stabilitv in the cerv
spill(:. Clill Orl/top, l(N, 85-96.
.
White, A.A. III & Panjabi. M.\1. (1990). Clillical Bio/llechal
oj' tlze Spille. Philadelphia: J.B. Lippincott.
Wittenberg, R.l-I., i\loeller, J.. &. \Vhite, A.A. III. (1990). C
pressive strength 01" autogenous and alJogenolls b
grafts 1"01' thoracolumbar and cervical spine fusion. Spi
15(101. 1073-1077.
Yang, K.H., Latour. B.K., & King, A.1. (1992). Computer s
ulation of occupant neck rtsponse to airbag deploymen
I"rontal impacts. J J3io/llech Ellg, 11.J, 327-331.
)'oganandan, N., Kumaresan, S.c., Voo, L., ct a!' (1996a).
nite dement modeling of dll: C4-C6 cervical spine u
lIed Ellg Phys, 18,569-574.
Yoganandan, N., Pintar, EA., \Lliman, D.J., ct al. (1996
I-Iuman head-neck biomechanics under axial tension. l
Ellg Pltys, 18,289-294.
)'00, J.U., Zou, D.. Edwards, W.T., et al. (1992). Effect of
vic,d spine motion on the netlroforaminal dimensions
human c('['vical spine. Spinc, 17{ 1131-1136.
"~~--------------~...",,.,...,..,..-----~~..... ....-. . -
._
ik
~
Biomechanics
of the Shoulder
~:
YJ,j
i>'
f/,
Craig 1. Della Va/Ie, Andrew S. Rokito, Maureen Gallagher Birdzel
Joseph D. Zuckerma
Introduction
Kinematics and Anatomy
Range of Motion of the Shoulder Complex
Sternoclavicular Joint
Acromioclavicular Joint
Clavicle
Glenohumeral Joint and Related Structures
Glenoid labrum
Joint Capsule
Glenohumeral and Coracohumeral ligaments
Additional Constraints to Glenohumeral Stability
Scapulothoracic Articulation
Spinal Contribution to Shoulder Motion
Kinetics
Muscular Anatomy
Integrated Muscular Activity of the Shoulder Complex
Forward Elevation
External Rotation
Internal Rotation
EXiension
Scapulotr,oracic Motion
loads at the Glenohumeral Joint
The Biomechanics of Pitching
Summary
References
It
----._------------------------~--------~._-._-~-------_._-
"""
:;;;Cor
Introduction
shoulder links the LIpper extremity to the trunk
acts in conjunction with the elbow to position
hand in space for crfkient function. [t consists
of the glenohumeral. acromioclavicular, stcrnoclav·
iculat: and scapulolhoracic articulations and the
musculature structures that act on them La produce
the most dynamic and mobile joint in the bodv (Fig,
12-J). An absence of bony constraints allo\I.'5 a wide
range of motion at the expense of stability, which is profor instead by the various ligamentous and musstl1.!cturcs.
The biomechanics of the shoulder arc complex,
and a complete discllssion necessitates an analysis
of the four aforementioned articulations that make
up the shoulc!eI- complex. This chapter describes the
Acromioclavicular
iOin1
@
~
/.
I
/'
\
,
'\
Sternum
_ r
($
I \
t?r
~,\
i"-(l--
,,
,
\
,
Humerus
I
~0L-n
Scapulothoraclc
I
i
,
I
I
,I
J.I
To produce the intricate motions necessary for no
n1al functioning of the shoulder complex, the fou
aniculations with their associated components ac
togcther in a way that produces mobility greate
than that afforded by anyone individual aniClda
tion, The ability of the shoulder complex to positio
the humerus and the remainder of the LIpper cx
trcmit~· in space is further augmented by movemen
of the spine. A discussion follows of the t\'pcs an
ranges of motion for the shoulder complex as
whole, with subsequent sections discussing th
manner in which motion is achieved at each of th
articulat ions,
RANGE OF MOTION OF THE SHOULDER
COMPLEX
7'
Acromion o f .
scapula
\ Clavicle
\
,)" I
Kinematics and Anatomy
Stern1i~:~tCular
-~
Scapula
anatomy of the various aspeclS or the shoulder com
plex and shows how their structure allows ror eff
cient biotllcchanical function.
_
Schematic depiction of the bony structures of the shoulder
and their four articulations. The circular insets show front
views of the three synovial joints-sternoclavicular,
acromioclavicular, and glenohumeral-and a lateral view
of the scapulothoracie joint, a bonemuscle-bone articulation. Adapted with permis5ion from De
Palma, A.F. (1983). Biomechallin of the shoulder. In Surgery of
the Shoulder (3rd 00.. pp.6S-85J. P!I;/adelp!l;d: J.8. Upp;ncorr
Shoulder range of motion is traditionally measure
in terms of flexion and extension (elevation o
movcrncnt of the humenls away from the side of th
thorax in the sagittal plane), abduction (elevation i
the coronal plane), and intemal·external rotalio
(axial rotation of the humerus with the arm held i
an aclducted position) (Fig, 12-2), Although durin
functional activit.ies these pure motions are rarel
seen, we can better understand the complex mo
tions of the shoulder by anal~/zing the separate com
ponents needed to achieve anyone position,
Although forward elevation of 180° is theOl'et
cally possible, the average value in men is 167° an
in women it is 171°. Extension or posterior eleva
lion averages 60· (Boone & Azen. 1979). These va
ues are Iimiled by capsular torsion. Abduction i
the coronal plane is limited by bony impingemen
of the greater tubcr()~ity on the acromion, Forwar
elevation in thc plane or the scapula, thereFore,
considered to be more Functional because in th
plane the inferior portion of the capsule is no
twisted and the musculature of the shoulder is op
timally aligned for elevation of the arlll (Fig. 12-3
Although shoulder range of motion !10nl1all~' d
creases as part of the aging process, physical activ
ity can counteract lhis process (Murray ct al
1985).
31
"~ ~'~"~T" "'-_" "_' O' ' ' ' ' =' ' 'r. ....=" "~" ,.,. .,." ,- , ,_: " ' ' '-:' ' '-:' ' ~- ,' ' '., ,~, "'.=.". .- -". . . ,.-. .,.-"_"', ."'_,"', _J"'.;./"', ~. -7"+-r;" ':, :-o"'- :;'?"'~ _P ~"'"'
._--Ii!_'.,
_,..
•.=..
.. T":::;"': ..
Extension
Forward flexion
~ it
lf
,
I
,
I
,
I
I
I
,
,,
:,
I
I
,,
,,
:,
--c:+
cs. ..- _........"..'...........-=>1
f
I
,,
I
I
!:::xternal and Internal
rolation
AbduClion
Humerus
at the
side
\\
":
'1,1
c
B
A
A, Forward flexion. The humerus is in the sagittal plane. B,
Extension. The humerus is in the sagittal plane. C, Abduction.
The humerus is in the frontal plane. D, Rotation around the
long axis of the humerus. External and internal rotation with
STERNOCLAVICULAR JOINT
The sternoclavicular joint consists 01" the enlarged
medial end of the clavicle and the most superolateral aspect of the manubrium, linking the upper extremity directly to the thorax. In addition, a small
facet is present inFeriorly that articulates with the
Scapular plane elevation
Long axis
Long axis
c3~~
Elevation in the scapular plane, which is midway between
forward flexion and abduction. The humerus is in the
plane of the scapula.
D
the humerus at the side, Internal rotation is shown with
arm behind the back, which is a functionally important
of this motion.
first rib. This is a truc s\TH)vial joint that has a
dle-like shape and contains a fibrocartilaginou
ticular disc or meniscus that divides it into
Cotlllxlrtmcnts.
Although the joint itself has little intrinsic st
ity, the articular disc in conjunction with ante
posterior. C()stoclavicular, and intL't"clavicular
ments maintains joint apposition (Fig. 12-4).
anterior and posterior sternocla\'icular ligam
resist anterior ;:md posterior translatiolls (as we
superior displacement), while the costoclm'i
lig;:Hllenl, which rUIlS betwcen the undersurfac
the medial end 01" the chlviclc and the first rib
sists upward as Wi..'l1 as posterior displacemen
the clavicle (\'ia its anterior'portion) and is tho
to be the major constraint in limiting stcrnoclav
br motion. The interclavicular ligament conn
the superollledial aspect or the clavicles ,:Hld as
in restr;:lining the joint superiorl:v. '1'he post
portion 01" the interclavicular ligament also. as
with anterior restraint of' the sternoclavicular j
Speciflcall)-', the interclavicular ligament tigh
with arm depression and is lax when the ann is
vated (i\'lorrey & An, 1990). The disc prevents
dial disphlcelnent
the clavicle, which m<l!' o
when C<\IT)-'ing objects at the side, ;:IS well as infe
or
Sternoclavicular Joint
Ant. sternoclavicular
Jig.
Articular
disc
ally directed coroid and the anterolaterally d
rected trapezoid ligaments, which arc distin
structures that serve different biomechanic
functions. The smaller coroid ligament acts
limit superior-inferior displacement of the clav
cle. The quadrilaterally shaped trapezoid is t
larger and stronger of the two ligaments and
found lateral to the coroid; it resists axial com
pression or motion about a horizontal axis. T
Post.
Protraction-retraction
Demonstration of the anatomy of the sternoclavicular
joint. Reprinted with permission from DeLee J. & Drez, D.
I
p~
(/9911). Orthopaedic Sports Medicine. Principles and Practice
(p. 464). Phi/adelfJhl~): WB. Saunders Co.
Sternum
A
displacement via articular contact. Although these
structures act as important stabilizers, they still al~
Imv for significant motion including up to 50° of axial rotation and 35° of both superior-inferior cleva·
lion and anteric)I'-posterior translation (Fig. 12-5).
Elevation-depression
Clavicle
Meniscus
l:7==
.......
ACROMIOCLAVICULAR JOINT
The acromioclavicular joint (Fig. 12·6) lies between
the lateral end of the clavicle and the acromion of
the scapula (the lateral and anterior extension of the
scapular spine) and is subject to the high loads
transmitted from the chest musculature to the upper extremity. It too is a synovial joint, but has a planar configuration. A wedge-shaped articular disc,
whose function is poorly understood, is found
\vithin the joint originating from the superior aspect. Both sides of the articular surface are covered
with fibrocartilage and the joint itself slopes inferomedially, causing the lateral end of the clavicle to
slightly override the acromion.
A weak fibrous capsule encloses the joint and is
reinforced superiorly' by' the acromioclavicular ligament. The acromioclavicular ligament acts primarily to restrain both axial rotation and posterior
translation of the clavicle. The majority of the
joint's vertical stability is provided by the coracoclavicular ligaments that suspend the scapula
from the clavicle (Fukuda et aI., 1986). The coracoclavicular ligaments consist of the posteromedi-
\
Ant.
•
B
Rotation
c
Motion at the sternoclavicular joint. A, Top view showin
the clavicular protraction and retraction (anteroposterio
gliding) in the transverse plane around a longitudinal ax
(solid dot) through the costoclavicular ligament, not
shown. Motion takes place between the sternum and th
meniscus. B. Anterior view showing clavicular elevation
and depression (superoinferior gliding) in the frontal pla
around a sagittal axis (solid dot) through the costoclavic
lar ligament. not shown. Motion occurs between the cla
cle and the meniscus. C, Anterior view depicting the clav
ular rotation around the longitudinal axis
the clavicle
of
•
Coroid ligament
Trapezoid
li)7ment
Coracoclavicular
ligamenl
Acromioclavicular
ligament
- Acromion
Coracoacromial
Intertubercular
synovial sheath -1'-----')
vasclll~ll'
slrlH.:lllrl"S and SL'IYL'S as an altachmcn
for Inan\ of Iht.' l1luscles dl41t ~Ict on tilL' sho
Tht.' clavicle' also pro\'ides for the normal ap
ance 4\[H.! contour or tIJL~ llPP(~J' chest. Elevati
IhL' upper extrcmil,\' is aCCOl11lXlnicd b.\· rotati
wdl as ck'\-',<1lion or the c1a\'idL.', with approxim
-1-0 of c1m·icular dL.'\·ation for evcr~· 10'> of arm
\·~ltion, \\·ith the majorit.\, of this motion occu
at the stL'I"llocla\'indar joint {Inman, Saunde
Ahhot!. 1944 J.
GLENOHUMERAL JOINT AND RELATED
STRUCTURES
Humerus
Scapula
The coracoclavicular ligament complex consists of the larger
and heavier trapezoid ligament, which is oriented laterally,
and the smaller coroid ligament, situated more medially.
Reproduced with permission from Hollinshead, WH. (/969).
Anatomy for Surgeons (Vol. 3), New York: Harper & RO'll,
coracoacromial ligament lies on the lateral side of
the acromioclavicular joint and runs rrom the
most lateral aspect of the coracoid to the medial
aspect of the acromion.
Although clavicular rotation does occur with arm
elevation, Rockwood (1975) found little relative motion between the clavicle and acromion. This has
been attributed to synchronous clavicular and
scapular rotation with little resullant relative motion at the acromioclavicular joint: the majority of
scapulothoracic motion occurs via the stcnlOclavicular joint. Thus, rigid fixation or fusion of the
acromioclavicular joint produces little loss or overall shoulder function while sternoclavicular fixation
leads to restricted motion or, more commonly. implant failure.
Although motion at the acrol'nioclavicular, st
chlvicular, and scapulothoracic articulalions is
10 the on:rall function or Ihe shoulder comple
cClltral pla.n?r is the glt.'llohullleraljoinl. Tht..' a
lar surface of lhe proximal humerus forms a
arc and is CO\·cfL.,d with hyalin(' cartilag~ as
glenoid fossa. The hUllleral head is rctnwt..'rlc
postt..'riorl.\' dirt..'ctcd 3D'" wilh rt..'speci 10 tilL>
cond.\'lar plan..: of the distal hUlllerus and has a
ward or llledial indination of 45°; this configur
gives the humerus an o\'eral! more anterior an
eral orienl~\liol\ (Fig. 12-7). The greater and
tuberosities lie lateral to Ihe articular surface o
proximal hUIl1L'rus thal SCIY('S as tht.' all<:lchlllL'1
for IhL' rotator culT IllUsctdalUrc. The long he
Lhe biceps lendon trii\vcrsL.'S till> bicipilal g
CLAVICLE
The clavicle lies between the two aforementioned
articulations, acting as a strut connecting the thorax to the upper extrell1it~'. It is an S-shaped,
double-curved bone: the medial two thirds of the
body convex anteriorly while the lateral end is concave. It protects the lInderl~,..ing brachial plexus and
The two-dimensional orientation of the articular sur
the humerus with respect to the bicondyrar axis, Rep
......rth pemllC;SIOfl from Rockwood, C. & Matsen, f (J 990).
Shoulder (p. 219). Pillladelphia. ~'V.B. 5.:1lmderc; Co
lies between the tuberosities) beneath the
tr",nsve,,'se humeral ligament (Fig. 12~8).
The proximal humerus articulates with the glenfossa, which itselF is retroverted 7° and superi~
inclined 5° relative to the plane of the scapula
12-9, A & B). This slight superior inclination
n,-OV\C!,'5 a significant degree of geometric stability,
to resist inferior subluxation or dislocation
et al., 1992). The glenoid fossa is shallow and is
to contain only approximately one third of the
di,wlet"r of humeral head. The bony architecture is
by the cartilaginous surface, which is
thicker peripherally than it is centrally, acting to
slightly but significantly increase the depth of the
Greater
tuberosity
Transverse
humeral
ligament
----1''';..-1
---+
Long head of - - - t the biceps
Short head of -----,f-tf-the biceps
I
B
A
A, The glenoid is retroverted r with respect to the p
perpendicular to the scapular plane. B, The glenoid f
superiorly approximately 5°. Reprinted ,·vith permission
from Simon, S,R, (Ed). (1994). Orthopaedic Basic Science
(p. 526-527). Rosemont. IL: AtlQS
glenoid as a whole. Although the congruity betw
the articular surfaces of the proxin1al humerus
glenoid was previously' thought to be somewhat
precise, stereophotogrammetric studies (\vhere
fine grid is projected onto an object that is then p
tographed From two different directions in the
text of known reference targets, allowing for a th
dimensional reconstruction) have shown
articulation to be precise, \vith the deviation f
sphericity' of the convex humeral articular sur
and concave glenoid articular surface being
than 1% (Soslowsky et aI., 1992). Less than 1.5
of translation of the humeral head on the gle
surface has been demonstrated in normal subj
during a 30 arc of 1110tion (Poppen & Walker, 19
thus, motion at the glenohumeral joint is alm
purely rotational. Given the paucity 01' bony
straint, stability is instead provided by the caps
ligamentous, and muscular structures that
round the glenohumeral joint.
0
The long head of the biceps goes in the bicipital groove
between the greater and lesser tuberosities. The transverse
humeral ligament helps stabilize the biceps tendon in the
groove. Reprinted with permission from Neer, C. (1990). Shoulder Reconstruction (p. 29), Philadelphia: WB. Saunders Co.
Glenoid Labrum
The glenoid labru111 (Fig. 12-10) is a fibrocarti
nous rim that acts to deepen the glenoid, provi
50°;(; of the overall depth of the glenohumeral j
(\Varner, 1993). It has a triangular configura
\'.'hen viewed in cross-section and has firm att
Coracoacromial ligament
Glenoid --',--'\l;--i.
labrum
Coracoid
process
Glenoid --+--~:
cavity
Coracohumeral
ligamenl
Tendon of long head of
biceps brachii
Long head of
tTiceps brachll
A
A, The glenoid labrum is attached to the underlying bony
glenoid and is confluent at its area with the long head of the
biceps tendon. B. The labrum has a triangular configuration
menls inferiorly to the underlying bone, having
morc variable and looser atlachments in its superior
and anterosuperior portions. The superior portion
or the glenoid labrum is connucnt with the tendon
of the long head of the biceps and. along with the
adjacent slIpraglenoid tubercle. serves as its site of
insertion (Moore. 1999).
iVlcasurcmenlS of the force needed to dislocate
the humeral head under constant compressive
pressure have shown that with an intact labrum,
the humeral head resists tangential forces of approximately 60% of the compressive load; resection of the labrum reduces the effectiveness of
compression·stabilization by 20% (Lippitt ct aI.,
1993). Detachment of the superior labrum with anterior-posterior extension (i.e., "SLAP lesion") can
occur from traction (repetitive overhead activities
or a sudden pull on the arm) or compression (a fall
onto an outstretched arm). This lesion can be a
cause of severe pain and shoulder instability (hoi,
Hsu, & An, 1996) as a result of significant increases in glenohumeral translation whcn compared with the intact shoulder (Pagnani et al.,
1995). (Fig. 12-11).
when vie'Ned in cross·section and serves to effectively
deepen the glenoid, increasing the stability of the glen
hllmeral joint.
DEmll.-
Oblique coronal image showing a tear involving th
sertion of the long head of the biceps tendon and
superior labrllm (arrow).
~----------
Capsule
'/fZ;/-:- The ~:dcnohLll11cral joint capsule has a significant de-of inherent laxity with a stll-face area llull is
$: -~~~~t\l"rec
,,,,,,"'-~"-;:;>
<"W'~""twice that of the humeral head (Warner, 1993), This
::~jZkJLredtindancyallows for a wide range of mOL ion. Mc,;,;" "idiallv, the capsule attaches both directly onto (an:~;~:;.:terotnferior1y) and beyond the glenoid labrum and
~~-- laterally it reaches [Q the anatomical neck of the
;'~lh~humerlls. Superiorly, it is attached at the base of the
coracoid. enveloping the long head of the biceps
~'7F~;~tendon and mnking it an intl'n-articular structure
dfB"HFig, 12-1 I),
-!:%ihi The capsule nlso has a stabilizing role, tightening
,-._. with various arm positions. In adduction, the capsule is Lattl superiorly and lax inferiorly; with abduction of the upper extremity, this relationship is
,'ev'er<;ect and the inferior capsule tightens. As the
is externally rotated, the anterior capsule tightwhile internal rotation induces tightening pasteriOl'lv, The posterior capsule in particular has been
to be crucial in maintaining glenohumeral
stability, acting as a secondary restraint to anterior
dislocation (particularly! in positions of abduction)
as well as acting as a primar~y posterior stabilizing
structure (Itoi, Hsu, & An, 1996),
'itrtY:'
Glenohumeral and Coracohumeral
ligaments
The three glenohumeral ligaments (SUPCriOI~ middle, and inferior) are discrete extensions of the anterior glenohumeral joint capsule and al'e critical to
shoulder stability and function (Fig, 12-12), The superior glenohumeral ligament originates from the
anterosupcrior labrum, just anterior to the long
head of the biceps, and inserts onto the lesser
tuberoSity. It is present in the majority of shoulders
bllt only well developed in 50%, The superior glenohumeral ligament acts as the main restraint to inferior translation with the arm in the resting or addllcted position (Warner el aI., 1992),
The coracohumeral ligament originates from the
lateral side of the base of the coracoid to insert on
the anatomical neck of the humerus (Fig, 12-6)
(Cooper et aI., 1993). This structure lies anterior to
the superior glenohumeral ligament and reinforces
the superior aspect of the joint capsule, These ligaments span the rotator interval belween the subscapularis and supraspinatus, Some research incHcates thal these strUClures have a secondary role in
preventing 'inferior translation of the shoulder while
in the adducted, neutrall.v rotated posilion. The
functional significance of the coracohumeral lig
rne-nt, however, seems to be related to the overa
development of the glenohumeral ligaments in
given individual, having a larger role in those with
less-developed superior glenohumeral ligame
(Warner et aI., 1992),
The middle glenohumeral ligament originates i
ferior to the superior glenohumeral ligament (at th
I o'clock to 3 o'clock position, right shoulder) an
inserts fl.1rther laterally on the lesser tuberosit
Great variability in the anatomy of this structu
has been demonstrated, being absent in as many
30% of shoulders (Curl & Warren, 1996), It ma
originate from the anterosuperior portion of th
labrum, the supraglenoid lubercle, or the scapul
neck, Morphological variants have been descl'ibe
including a cord-like variant (clearly distinct fro
the anterior band of the inferior glenohumeral lig
ment) and a sheet-like variant (blending with th
anterior band of the lnferior glenohumeral lig
ment). Functionally, the middle glenohumeral lig
ment acts as a secondary restraint to inferior tran
lations of the glenohumeral joint with the arm
the abducted and externally rotated positio
(vVarner el aI., 1992). It also sen/cs as a restraint
anterior translation, having its maximal effect wi
the arm abducted 45°.
The inferior glenohumeral ligament originat
from the inferior aspect of the labrum and inser
on the anatomical neck of the humerus. It has bee
shown to have three distinct components (O'Brie
et ai., 1990): an anterior band originating from 2
4 o'clock (right shoulder), a posterior compone
originating fTom 7 to 9 o'clock (right shoulder), an
an axillat), pouch (Fig. 12-12), The inferiOl· glen
humeral ligament has the greatest rll11ctional signi
icance, acting as the primm)' anterio.' stabilizer
the shoulder with the arm in 90' of abduction, A
the arm is abducted and externally rotated. the a
terior band of the inferior glenohumeral ligame
tightens, resisting anterior tninslalion. \·Vith inte
nal rotation of the abducted arm, the posterior ban
becomes taut and posterior translation is resiste
The inferior gleno·humeral ligament complex al
serves to resist inferior translation of the glen
humeral joint with the arm in the abducted pos
tion. Variability as to the size and attachnlent sit
of the gle-nohumeral ligaments has been demo
strated (Warner et aI., 1992); howe vet; the clinic
significance of this has yet to be fully elucidated a
though it has been suggested thal absence of t
middle glenohumeral ligament may predispose
instability (Steinbeck et aI., 1998),
A-
B
A
)
AS
c
D
A, Schematic drawing of the shoulder capsule illustrating the
location and extent of the inferior glenohumeral ligament
complex (IGHLC). A, anterior; P, posterior; B, biceps tendon;
SGHL, superior glenohumeral ligament; MGHL, middle glenoo
humeral ligament; AS, anterior band; Ap, axillary pouch; PS,
posterior band; Pc, posterior capsule. Reprinted with pertnis·
sian from O'Brien, SJ, Neves,
M.e,
Amoczky, S.P, et
at.
(1990)
The anatomy and histology of the inferior glenohumeral ligament
complex of the shoulder. Am J Sports Med. 18, 579-584. B. The
superior glenohumeral ligament is the primary restraint to
inferior translation in the adducted shoulder at neutral rotation. In this position, the middle glen?humeral ligament and
the anterior and posterior bands of the inferior gleno·
humeral ligament complex remain lax, C, The anterior b
is the primary restraint resisting inferior translation of t
shoulder at 45<> abduction and neutral rotation, In this
tion, the superior glenohumeral ligament, the middle g
humeral ligament, and posterior band are lax, D, At 90°
duction, the anterior and posterior bands of the inferio
glenohumeral ligament cradle the humeral head to pre
inferior translation, The posterior band is more significa
external rotation, whereas the anterior band plays a gr
role in internal rotation. Reprinted wirh permission from W
IP, Deng, X.H.. Warren, R,F., et al. (1992). Staric capsuloliga
taus restraints ro superior-inferior translations of rhe glenohu
joint. Am J Sports Med, 20, 675-678.
Experimental Techniques:
Ligament Cutting Studies
Shoulder Instability
A
Ligament cutting studies have been instrumental in fur-
.
thering our knowledge regarding the contribution of a
21-yea:"0Id ma.n fell while skiing onto his right upper
extremity, cauSing forceful abduction and external ro"
given anatomical structure to overall glenohumeral sta-
tation. He noted acute pain in the arm and was unable to
bility (Curl & \JVarren, 1996). In this technique, cadaveric specimens are biomechanically tested before and
IYlQveit Physical examination revealed loss of the normal
contour of the shoulder and painful range of motion. Ra"
dlogr~phs showed an anterior dislocation with posterior
after selectively cutting sequential structures. A force is
then applied in a given arm position, and the transla-
superi.()rhumeral head impaction fracture. The patient un-
tion that occurs is measured. From this information, the
gerwent closed reduction. Postreduction radiographs con"
relative contribution that a given structure provides to
firlYledreduction of the humeral head with a small bony
overall stability can be determined. When a particular
ayulsion fracture of the anterior-inferior glenoid rim,
pattern of shoulder instability is then identified, the
whichrepresents a detachment of the anterior labrum in
physician can infer which anatomical structures may be
the area of the superior band of the glenohumeral liga-
deficient or disrupted so that a rational plan of physical
ment insertion (Case Study Figure 12" 1" 1).
Structural alteration of bony geometry, ligaments, and
therapy or surgical repair can be implemented.
•
labrum resulted in shoulder instability. The detachment of
the anterior labrum and the superior band of the ligament insertion affected primarily the resistance to anterior
Additional Constraints to Glenohumeral
Stability
Sy'l1ovial fluid acts via cohesion and adhesion to
further stabilize the glenohumeral joint. Synovial
fluid adheres to the articular cartila~e ovcrh;in cr the
glenoid and proximal humerus, c;using the b two
surfaces to slide along one another: Th~ synovial
fluid provides a cohesive force bet\veen the~e two,
making it difficult to pull them apart (Simon,
1994). Under normal conditions, the intra-articular
pressure \vithin the gknohumeral joint is negative,
acting to pull the overlying capsule and glenohumeral ligan1cnts inward. If the integrity of the
glcnohun1cral joint capsule is cornpro~1is~d (e.g.,
venting the capsule) or if a significant effusion exists (normally the glenohumeral joint contains less
than 1 cc of fluid), significant increases in translation are observed (Kumar & Balasubramaniam,
1985). Specifically, venting the capsule reduces the
force needed for anterior humeral head translation
by 55 0ft;, for posterior translation bv 430c), and for
inferior translation by 57% (Gibb et ·al., 1991) (Case
Study 12-1).
translation of the humeral head, resulting in anterior dislocation. !n addition, a concomitant capsule lesion affected the intra"articular negative pressure necessary to
pull the humeral head inward. After conservative management. the patient did well for 6 months until he sus"
tained a recurrent dislocation. The patient subsequently
opted for operative treatment that included repair of the
fracture and capsule and a complete period of rehabilita"
tion~vithernphasis on joint stability and proprioception.
SCAPULOTHORACIC ARTICULATION
The scapula is a Flat, triangular bone that lies on
the posterolateral aspect of the thorax between the
second and seventh ribs. It is angled 30° anterior to
the coronal plane of the thOl;X and is rotated
slightly toward the midline at its superior end and
tilted anteriorly with respect to the sagittal plane
Case Study Figure 12-1"1.
(Fig. 12-13) (Laumann. 1987). The spine of the
scapula gives rise laterally to the acromion process
that articulates with the distal clavicle at the
acromioclavicular joint. The coracoclavicular ligaments and muscular attachments help to suppon
the scapula and stabilize it against the thorax (Fig.
12-6). There is, however, no osseous connection
with the axial skeleton. This allows for a wide
range or scapular Illotion, including protraction,
I"etraction, elevation, depression, and rotation.
The scapulothoracic articulation involves gliding
of the scapula on lite posterior aspect of the thorax.
Interposed between the scapula and the thoracic
\vall lie the subscapularis (arising from the costal
surface of the body of the scapula) and the serratus
anterior, which help to stabilize the scapula against
the chest wall and thus prevent "scapular winging"
(Fig. 12-14). These two muscles glide along one another to provide greatly enhanced mobility of the
shoulder complex as a whole,
Elevation of the arm involves rnotion at both the
glenohumeral joint and the scapulothoracic articulation. Although the contribution from each varies
according to arm position and the specific task bclng performed. the average ratio of glenohumeral to
scapulothoracic moL ion is 2: I (Tibone el aI., 1994).
Elevation of the arm also induces complex rotatory
motion of the scapula, \vith anterior rotation during
the first 90° followed by posterior rotation with a total arc of approximately 15° (Morrey & An, 1990).
-.J 3 'r Scapular
I
I
Anterior view of the scapulothoracic articulation, a
I11llscle·bone articlJlation between the scapula and
During scupular motion, the subscaplJlaris muscle. Io
attaches broadly to the costa! surface of the scapula
on the serratus anterior muscle, which originates or
first eight ribs and inserts into the cost<ll surf..'!ce of
scapula along the length of its vertebral border.
.---------------SPINAL CONTRIBUTION TO SHOULDER
MOTION
Allhaugh often O\'crlookcd, motion 01" the th
and lurnlmr spine contributes to thL' :.\bilit:,· t
tion the upper c."tl'(.'l1lit~· in space, {ht.T('b~· l.'
ing the on.:r<.111 Illotion and functioll of the sh
cOlllpk.·x. f'kxiol1 01" tile spine away frolll an ,il~' altempting ttl reach an object o\'crhcacl t.'n
the range of motion <.lllainahlt: (Fig. 12·15), T
ponancc or spinal motion in O\·l..'rlH.':'ld act
such as throwing ~Illd r<.tcqul.'l sports h<.ts als
c1emonst rated.
Acromion
0
~l
SerraIus
anterior
spine
I.
Scapula
Kinetics
Scapular orientation on the chest wall. Left.
30~
anterior,
Right. 3" upward, Reprinted with permission from Warner JJP:
The gross anatomy of the joim surfaces, ligaments, labrum, and
capsule. In: Matsen. F.A., FtJ, F.H., Hawkins, R.I (Eds.) The Shoul-
der: A Balance of Mobility and Stability. Rosemont, IL: American
Academy of Orthopaedic Surgeons, J 993.
Numerous llluscll..'s aCI nn the \'arious comp
the shoulder complex to provide both m
and d\'namic $t~lbilit.\·. Dynamic stabilizati
cllrs via sl.~\'cral possible lllcch:'lnisms (Mo
An, 1990), including p~lssi\'(' musch.' lensi
via a b:,uTic:r effect of the contracted Illuscle
prl..'ssi,·L' forces broughl ahout h~' llHISCU!:'U
traction, .ioint lllotion that induces lighten
or
"fmm'-----I
activated differently for specific activities. The
toid originates from the lateral third of the clav
acromion, and scapular spine and inserts on the
terolateral aspect of the humerus. The anterior h
acts as a strong flexor and internal rotator of
humerus, the middle head as an abductor, and
posterior head as an extensor and external rot
The pectoralis major lies over the anterior c
wall and has two heads, a clavicular head orig
ing from the side of the clavicle and a sternoco
head originating from the sternum, manubri
and the upper costal cartilages. The two heads
verge at the sternoclavicular joint. The pecto
major inserts at the intertubercular groove of
humerus between the tuberosities and act
adduct and internally rotate the humerus. Se
darily, the clavicular head acts as a flexor or fon
elevator of the humerus while the sternocostal
extends the humerus. The pectoralis minor lies
to the pectoralis major, functioning as an impor
scapular stabilize!: The pennate subclavius mu
lies inferior to the clavicle and Illay assist in cl
ular motions. It has a tendinous origin from the
teromedial aspect of the first rib and insert
lateral bending of the spine enhances the ability to position the upper extremity. Reprinted with permission from
Rockwood, C. & Matsen, F (1990), The Shoulder (p. 2' 9).
Philadelphia: WB. Saunders Co.
Pectoralis major
Clavicular
Sternal
Deltoid
Posterior
Middle
Anterior
the passive or ligamentous constraints, or via a
redirection of the joint force toward the center of
the glenoid.
To understand rnusclc function and force transmission. one must consider a given muscle's orientation, size, and activity. Given the multiple articulations present in the shoulder complex, an:y given
muscle may" span several different joints, and depending on the position of the upper extremity', its
relationship with regard to anyone articulation
may change, altering its effect on that joint and the
resultant forces or motions produced.
MUSCULAR ANATOMY
The shoulder musculature can be thought of in layers. The outermost la)-!er consists of the deltoid and
pectoralis major muscles (Fig. 12-16). The deltoid
forms the normal, rounded contour of the shoulder
and is triangular in shape, with anterior, middle,
and posterior heads. Each portion of the deltoid is
Serra
ante
Anterior sheath
of rectus
Pectoralis
minor
Anterior view showing the superficial muscles (left sh
der) and the deep muscles beneath the deltoid and p
toralis muscles (right shoulder).
the undersurface of the medial clavicle (Morrev &
Anterior View
An, 1990).
BcnL~alh this ouler layer lies the rotator cuff
musculature: the supraspinatlls. infraspinatus.
subscapularis, and teres minor (Fig.12-17). These
four muscles act to abduct and rotate the
humerus and act as important glenohumeral stabilizers via both passive muscle tension and dynamic contraction. The supraspinatlls originates
from the supraspinatus fossa of the scapula and
inserts on the greater tuberosity of the prox.imal
humerus. It forms a force couple with the deltoid
during abduction of the humerus. The infraspinalllS
Biceps
and teres minOt- originate from the inferior as-
pect of the scapula and insert on the greater
(uberosity. These muscles act as external rotators
of the humerus. The subscapularis lies on the
costal surface of the scapula and inserts on the
lesser tuberosity of the proximal humerus. It
functions as an important internal rotator of the
humerus. The subscapularis. along with tlte mid~
dIe and inferior glenohumcral ligaments, has also
been shown to act as an important anterior stabilizer of the glenohumcral joint, particularly with
lJ1(.> arm held at 45° of abduction. The teres major
muscle (Fig. 12-18), while not part of the rotator
culT, also originates from the scapula, but at its
inferior angle coursing inferior to the teres minor
and then passing anteriorly to insert on the
humerus at the intertubercular groove. h functions to assist with arm adduction and internal
rotation.
The biceps muscle is also involved with motion of
the shoulcleI· complex. It is composed of two heads:
a short head that originates fTom the tip of the coracoid process of the scapula and a long head that
originates I'T om the superior glenoid labrum and
supraglenoid tubercle (Fig. 12-8). The tendon of the
long head of the biceps lies within the glenohumeral
joint and descends between the greater and lesser
tuberosities. joining the short head to insert on the
bicipital tuberositv of the radius. The biceps fl1nclions to Oex and supinate the forearm and elevates
the humerus. The long head of the biceps also acts
as a humeral head depressor and, as such, plays a
role in maintaining glenohumeral stability (hoi et
aI., 1994).
Several muscles lie on the back and act directly
on the scapula (Fig. 12-18). The outermost layer
consists of the trapezium that covers the posterior
neck and uppermost portion of the trunk, inserting
on the superior aspect of the lateral one third of the
clavicle, acromion, and scapular spine. The tl'apez-
Subscapularis
A
Posterior View
Supraspinatus
"
Infraspinatlls
B
A, Anterior view. The "rotator interval" is a term
duced in 1970 to indicate the space between the
supraspinatus and subscapularis tendons. The cor
humeral ligament lies superficially along its ante
where it is readily available for release as indicat
long head of the biceps lies deep along its poste
and serves as a guide to this interval during surg
Reprinted ~'llth permission from Neer. C. (1990). Shou
construetiofl (D. 29). Pllild(/e~ohia: !NB. Saunders (0_
rior view_ The two external rotators of the hume
..infraspinatus and teres minor muscles. which are
posterior wall of the rotator cuff. Note the medi
of the infraspinatus, which is often mistaken at s
the border between the infraspinatus and the te
Reprinred with permission [rom Rock\'-Ioocl, C. & Mat
(/990). The Shoulcier (p. 2) 9). Philadelphia: \>';1 B. Sau
,l
L
Deltoid
Anterior
~-'\-- InfraspInatus
Middle
'-=--+- Rhomboideus
major
Posterior
Teres minor
Teres major
)
Latissimus
dorsi
Posterior view shoWing the superficial muscles (left shoulder) and underlying muscles
(right shoulder).
ium serves to elevate. retract, and rotate the
scapula. The latissimus dorsi covers the inferior
portion of the back, inserting on the intertubercular
groove of the humerus. It acts to extend. adduct,
and internally rotale the humerus.
Below these muscles.lie the levator scapulae superiorly, which elevale and inferiorly rotate the scapula,
and the rhomboid muscles below, \vhich retract and
rotate the scapula. Both of these muscle groups nctLO
assist the serratus anterior (which lies laterally on the
chest and intercostal muscles inserting onto the me~
dial border of the anterior surface of the scapula) in
keeping the scapula fixed to the trunk.
INTEGRATED MUSCULAR ACTIVITY
OF THE SHOULDER COMPLEX
.' .;:.
Electromyography allows for the quantification of
muscular activity during dynamic conditions.
This permits insight into the level of muscular ac-
tivity but does not directly indicate forces gene
ated. A complete understanding of the latter
quires knowledge of lhe moment arm (measur
as the distance between the instantaneous ccnt
of rotation of thc joint and the distance of mu
cular pull) and the physiological cross~section
the involvcd muscle ('measured as the muscul
volume divided by its length). In the should
complex, each motion is associated with mov
ment at multiple articulations and the constan
changing relationships of the muscular origi
and insertions.
Given the paucity of osseous stabilil~' at t
glenohumeral joint, force generated by one mu
cle: (the primary agonist) requires the activati
of an antagonistic muscle 50 that a dislocati
force does not result (Simon, 1994).. The antag
nist usually accomplishes this via an eccentl
contraction whereby the mU5cfe is lengthen
while actively contracting or via the production
a neutralizing force of equal magnitude but in the
opposite direction. The relationship between two
such muscles is also referred 1O as a force couple
(Fig. 12-19). Around the glenohumeral joint. there
is a force couple in the coronal plane (between the
deltoid and the infedor portion of the rotator
curf) and in the transverse plane between the subscapularis muscle anteriorly and the posterior rotator Cliff Illusculature (the infraspinatus and
teres minor).
Relative Illotion is produced by an imbalance bet\veen the agonist and antagonist that produces
torque. The degree of torque and the resultant angular velocity produced is determined by the relative activation of two such muscles or muscle
groups. Resultant muscular forces arc determined
via an understanding of the cross-sectional area of
the activated muscles involved and their orientation
at the time of activation.
A
B
The deltoid and the oblique rotator cuff muscles (infraspinatus. subscapularis. and teres minor) combine 10 produce elevation of the upper extremity by means of <1 force couple
(two forces equal in magnitude but opposite in direction).
With the arm at the side (A), the directional force of the deltoid is upward and outward with respect to the humerus
while the force of the oblique rOlator cuff muscles is downward and inward, These two rotational forces can be resolved into their respective vertical and horizontal components. The horizontal force of the deltoid acting below the
(enter of rotation of the glenohumeral joint is opposite in
direction to the horizontal force of the oblique rotalOrs.
which is applied above the center of rolation. These forces
acting in opposite directions on either side of the center of
rotation produce a powerful force couple. as illustrated by
the arm signal (B), The vertical forces offset each other,
thereby stabilizing the humeral head on the glenoid and allowing elevation to take place. Adapted with permission
from Lucas, 0.8. (1973). Biomechanics of the shoulder joint.
Arch 5urg. 107.425.
Forward Elevation
TIll' most b~\~,dc motion
(If thl.· shoukkr compk.
\·oln.'s ck\'~lIioll of tIlL' arm ill the scapular p
This motion has hel'n studiL'd in depth via both
ll'omyograph}' and stcrcopllOlogrammctl-.v.
muscles of the shoulder girdk han: subsequ
heen grouped according: to relatin.' illlponancL"
regard to this rnotion. TIll' first grouping inc
the dcILoid <Spccificall~.. lhe anterior and m
heads). lrapeziulll (inf..:rior ponion), supraspin
and serratus ~lJ)leri()r (Simon, 10(4), The se
grouping consists uf' the Iniddlc portion o
trapeziulll. the infraspinatus. and thl..' long he
the biceps. r\ third group consists or Ilw pos
head or the ddlOid. thl.' c1~l\·iclilar Ill,:'~ld of tht.~
toralis major, and lht.' slIpi.:rior ponion or the tr
iUIll. The fourth and final group includes I hi.' st
head or the pectoralis rnajor. the latissirnlls d
and t Itt.' long !It.'ad or the t ricL'ps.
Thl.' inter-rt..'!i.lriol1ship bL'l\\·ct.'1l the mus
rorces in\·oh'l.'d in shoulder dl.'\'~llion W<lS firsl
ied b,\' Inman, who round that the deltoid
sllpr~ISpiJlatus \ulrk s}·nergisticall.,· ,,·hile th
mainder of till".' rotator cufllllusculalllrc provi
hUllleral dcpn:ssing forct.' 10 counter sllbluxati
lhl.:.' humeral !lead (Illll1~\I1. Saunders. & Ab
19-1-1). Thus. the vertic'l1h· orien'ed pull o[ th
(oid is offset h~' a net inferior rorce crcat(:d b~' th
fraspinatlls, suhscapularis, anc! teres millOI'.
EIt.:.'ctrom.,'ographic studies have shown tlwt
lhe supraspilwluS and deltoid ~\rc acti\'L' throug
the nlllgt..' of ann clev"ltioll. ThL' supraspinatus.
ever, is felt to h~l\'c a larger role in initiating a
tion. As the arm is progrcs~i\'d}" dL'\'~\ted from
side, the 1ll0!llCI\t arm of till".' deltoid improvl.
suiting in a larger force in I\.'lation to thc supras
tus (Fig. 12-20), The p<.:rccllwge 01" shearing or
cal force created by the deltoid likewise ck'crt
with increasing arnollllts or ahduction. The an
pull or the supraspinalUs'is more l'OIlSI~lJll a
pro:dmalcl.,· 7,=,°, acting not onl.\· to elcvate or a
the arm but also to compress the hurneral
within the glenoid. Thl.' remaining roW tor cull
cles pull at approxil1latd~' -15°, which is din;(:ll"
{'edod}; (slightl,'· highcr in the ter('s minor at 55
sulting in forces lh~lt l.·quall}' compress ;:111(1 de
the humeral head to maintain glcnohulllcral s
ity (Fig. 12-21).
Sl~\('ctive anesthetic block or tht' n:dllary
(and resulting dcILoid paralysb) dcnlOnsll·atl.'
forward elevation is possihle alb~it signinc
weakencd. Lik.:,\·isc. a supr~lscaplilar nerve
Subscapularis
Supraspinatus
\
ff,o':::;7'1:_u..p",ra::-s-,-p_in_a-tIU.S
-~------­
I
As the arm is abducted to 90", the direction of pull of the
deltoid approximates that of the supraspinatus. Therefore, patients with a large tear of the rotator cuff often
can actively maintain the arm abducted to 90° but may
not be capable of actively abducting to 90", Reprinted with
permission from Simon, S.R. (Ed). (/994) Orthopaedic Basic
Supraspinatus
Science (p. 527). Rosemont, IL: AAOS.
I
--I-:~~~~~~~
9
and the resultant supraspinatlls paralysis induced
has a similar eHeet. However, a block of both
nerves results in a loss of arm elevation (Colachis,
& Strohm, 1971; Howell ct a\., 1986) (Case Study
12-2).
\Vhen pure abduction is compared with pure forward elevation, the same basic relationships are
seen with the rotator cliff acting to stabilize the
glenohumeral joint while the deltoid provides the
necessary torque. Forward flexion results in activation of the anterior and middle deltoid (73 and 62%
activity, respectively') \vith stability provided mainly
by the supraspinatus. infraspinatus, and latissimus
dorsi, the lalter being particularly active (25°;0 activation) with forward flexion beyond 90°. Pure abduction requires similar muscular activity; however, the subscapularis shows increased activation
as it acts as the prime stabilizer via eccentric contraction.
Infraspinatus -
Teres minor
External Rotation
The primary external rotator of the humerus is the
infraspinatus. with significant contributions made
by the posterior head of the deltoid and the teres minOI: With any given amount of shoulder abduction,
electromyography reveals the prime external rotator
to be the infraspinatus. The subscapularis is similarly active but serves an antagonistic role as the
main stabilizer preventing anterior displacement of
The angle of pull of the subscapularis (top) is approximately
45°. The angle of pull of the infraspinatus (bottom) is also
approximately 45°, and the teres minor (bottom) is approximately 55°. These vectors result in nearly equal glenohumeral joint compression and humeral head depression. The
supraspinatus (center) is essentially horizontal in its orientation, resulting in compression of the glenohumeral joint.
•
----_._.__
._----_._~~----~-_._._-----~-
----._-
Subacromial Impingement Syndrome and Rotator Cuff Tear
·.
A
53-year-old right-hand-dominant woman who presents
right shoulder pain of 6 months' duration, The patient
noted increasing pain at night and with overhead activities of
daily living such as putting on a shirt and brushing her hair.
The pain is unresponsive to pain killers. Physical examination
revealed diffuse tenderness over the shoulder and deltoid
with.painful forward elevation above 60:3 and internal fOla-
still exert Its effects by Wc1Y of the "spans" of the bridge.
Thus. patients may have a "functional" rotator Cliff tear in
\-vhich they are s!ill able to periorm overhead activities. Ii the
tearing force presem at the t\:'IO suppons is great enough.
howe'/er, worsening of the tear will occur.
tion limited to her ipsilateral greater trochanter. She had a
positive Neer sign (painful forward ele'l<ltion between 60 and
120°) and a positive Hawkins sign (pain with passive abduction and. internal rotation). A subacromial impingement test
was positive with near complete relief of pain and restoration
of fOl\lvard elevation to 1500 but persistent \;veakness on resisted shoulder abduction. A diagnosis of subacromial impingement syndrome is made. The patient was initially managed with conservative treatment. At 5 weeks follow-up. the
patient had persistent pain and an MRI of her shoulder revealed a full thickness tear or the supraspinatus tendon. The
patient opted for operative management (Case Study Fig, 12-1-1).
Biomechanically, rotator cuff tears have been likened to a
suspens.ion bridge, whereby the free margin of the tear corre·
sponds to the cable and the residual atlachrnents correspond
to the supports at each end of the cable's span. This coniiguration allows the muscle that has been torn at its insertion to
the humeral head with external rotation. As the
amount of shoulder abduction is increased, the pos~
terior deltoid increases in cff-!cicncy as an accessory
external rotator
the humerus secondary to improvement of its moment ann.
or
Internal Rotation
Internal rotation of the shoulder is accomplished by
the subscapularis. sternal head of the pectoralis major, lHtissinlus dorsi, and teres major. The subscapularis is active during all phases of' internal rotation,
with decreased relativc activity seen with cxtremes
of abduction. 1n the samc \Va.v. activity of the sternal
head of the pectoralis major Hnd the latissimus
dorsi decreases with abduction. However. the posterior and middle heads of the dehoid compensate
with increased eccentric activity during internal rotation while the arm is abducted.
Extension
Extension of the upper extremity is accompl
by the postcrior and middk heads of the de
The supraspinatus and subscapularis are also
tillually active throughollt arm extension, res
forces via eccentric activil~« that would tend to c
anterior dislocation.
Scapulothoracic Motion
at the scapulothoracic articulation a
maintenance or deltoid tcnsion, allowing
maintain optimal power regardless of arm
lioll. \-Vith forward elevation at the arm,
scapula rotates, increasing stability at the g
humeral joint and decreasing the tendency 1'0
pingement or the rOlator cufe bcncmh
acromion (Fig. 12-22). A rotalional rorc~ CO
(two equal. noncollinear, parallel but opposite
Motion
forces) bet\veen the upper trapezium, levator
and upper serratus anterior with concontraction of the lo\ver trapeziurn and
cP'T"llllS anterior leads to scapular rotation that is
for full forward elevation (Fig. 12-23)
1994).
.u,","""o> AT THE GLENOHUMERAL JOINT
glenohumeral joint is considered to be a
load-bearing joint. Although calculations
the exact forces acting on it are challenging
thc large number of involved muscular
and possible positions attainablc, sevsin1plifying assumptions allow an estimaof the magnitude of these forces. A freediagram of a person holding the upper
held at 90' of abduction can be utias an example; it is assumed that only! the
ue"cm, muscle is active and that it acts at a distance of 3 cm from the center of rotation of the
humeral head. Three forces are then considere
the deltoid muscle force (D), the weight of th
arm (equivalent to 0.05 body weight [BW] actin
at the center of gravity of the limb, 3 cm), an
the joint-reactive force at the glenohuIl1eral joi
(1). The joint reactive force and the force of th
deltoid, being nearly! parallel, are considered
be a force couple and are of equal but opposi
magnitude, The force on the glenohumeral joi
\vhen holding the arm at 90° of abduction can b
estimated to be one-half of body weight (Calcul
tion Box 12-1, Case A). 11' a weight (W) of 2 kg
added (equivalent to 0,025 body weight of an SOmale) to the hand of the outstretched extremi
held at 90° of abduction, a sirnilar calculation ca
be made (Calculation Box 12-1, Case B).
Experimentally, loads at the glenohumeral joi
and the forces necessary! for arm elevation hav
been detcrmined. These forces have been found
be greatcst at 90° of elevation, with the deltoid for
equivalent to 8.2 times the weight of the arm an
Forward elevation or abduction 0-120 o of the arm requires synchronous rotation of the
scapula. Reprinted "'lith permission {rom Simon, S.R. (Ed). (1994). Orthopaedic Basic Science (p.
1
527-535). Rosemont, IL: AAOS.
- - - - - - -
Rotation of the scapula is produced by the synergistic contractions of the lower portion
of the serratus anterior and the lower trapezius, with the upper trapezius, levator scapulae, and upper serratus anterior. Reprinred with permission from Simon, 5.R. (Ed). (1994).
Orthopaedic Basic Science (p. 527-535). Rosemont. IL: AAOS
the joint reactive force equivalent to 10.2 times the
weight of the arm (Inman, Saunders, & Abbott,
1944). More recent investigations of these same
forces utilizing the assumption that the muscular
force is proportional to its area multiplied by its ac~
tivity as determined by electromyography have
yielded similar values. with a maximal joint reactive
force of 89% of body weight seen at 90 0 of elevation
in the scapular plane (Poppen, & Walke.; 1978).
THE BIOMECHANICS OF PITCHING
Pitching has been divided into five stages: wind up,
earl.y cocking. late cocking, acceleration, and follow-through (Tibone et 'II., 1994). The deltoid has
been found to be responsible for elevation and abduction of the humerus in the carly phases. followed by increased activation of the ratmar cuff
musculature in the late cocking phase aCling b
rotate the humerus and prevent anterior su
tion of the glenohumeral joint (Eames &
1978). Specifically, the supraspinatus acts in t
cocking phase to draw the hUllleral head low
glenoid, the infraspinatus and teres minor p
humeral head posteriorly, and the subscap
bOlh prevents excessive external rotation
humerus ancl contracts eccentrically to
stress on the anterior shoulder (Tibone et 'II.,
The importance of scapular (and thus glenoi
bilization has also been recognized, and the se
anterior has been shown to nre actively in t
cocking phase; this provides a stable platfo
humeral motion. Thus. coordinated. sequenti
vation of the shoulder musculature is needed
vent anterior subluxation of the glenohumera
and the overuse tendinitis that can ensue.
~~,;;~
-/j
,0;"0" .Calculation
of Reaction Forces
Estimates of the reaGlon force on the glenohumeral
joint are obtained with the use of simplifying assump:, lions (Poppen & Walker, 1978).
;;;
Case A. In this example, the arm is in 90° of abduction,
;~(:and it is assumed
Ii
that only the deltoid muscle is active. The
force produced through the tendon of the deltoid muscle (0)
:,~,acts at a distance of 3 em from the center of rotation of the
,,'joint (indicated by the hollow circle). The force produced by
~, the weight of the arm is estimated to be 0.05 times body
\.veight (BvV) and acts at a distance of 30 cm from the center
of rotalion. The reaClion force on the glenohumeral joint 0)
may be calculated with the use of the equilibrium equation
o is approximately one-half body weight. Because 0
are of equal magnitude; thus, the joint reaction force is also
approximately one-half body weigll!.
Case B. Similar calculations can be made to determine the
value for 0 when a weight equal to 0,025 times body weight
is held in the hand with the arm in 90" of abduction.
~M
= 0
(30 ern x .05 BW) + (60 em " .025 BW) (D x 3 em) ~ 0
D = (30 em " .05 BW) + (60 em " .025 BW)
30m
that states that for a body to be in moment equilibrium, the
5;Jm of the moments must equal zero. In this example, the
and J
are almost paraUel but oppOSite. they form a force couple and
D = 1 BW
moments acting clockwise are considered to be positive and
counterclockwise moments are considered to be negative.
'::
~M =
Once again. 0 and J are essentially equal and opposite.
forming a force couple. Thus. the joint reaction force is ap-
0
(30 em x .05BW) - (D " 3 em) = 0
proximately equal to body weight.
D = 30 em x .05 BW
3cm
D = 0.5 BW
3cm
3cm
I
Force 0 ......... ----i""X'Force J
I
I
II I
!
I,
,.ossw
;•
,
30cm
):
Calculation Box Figure 12-1-1.
&.............. . .
,
;.
_
·········~)~~=----=t;;;;_;;:;---r
!
30cm
.~,----::------->
60 em
Calculation Box Figure 12-1-2.
__ .-._ _.
Sumrnarv
F The shoulder consists of the glenohumeral,
ncromioclavicular. stcrnoclaviculat: and scapulothoracic articulations and the muscular structures that
act on them to produce the most mobile joint in the
body.
2 The sternoclavicular joint, which connects the
medial end of the clavicle to the manubrium. links
the uppe.· ex trernity to the thorax. The aniculnr disc
within the joint and the ligaments that surround
provide stability while allowing for significant rota
tion of the clavicle. lillie relative motion is seen be
tween the clavicle and acromion at the acromio
clavicular joint.
3 The glenohumeral joint. while known to be
precise articulation, is inherently unstable becaus
the glenoid fossa is shallow and is abl.e to contai
only approximately onc third of the diameter of th
humeral head. Stability is instead provided by th
capsular, musculal~ Llnd ligamentous structures that
surround it.
4 The three glenohumeral ligaments (superior,
middle. and inferior) are discrete extensions of Ihe
anterior glenohumeral joint capsule and are critical
to shoulder slability and function.
S" The inferior glenohumeral ligament has the
most functional significance (particularly the
anterior bane!), acting as the primary anterior
stabilizer of the shoulder when the arm is 'lb·
ducted 90'.
·6. The inlegrity of the shoulder capsule and the
negative intra-articular force it maintains also plays
an integral pan in mainlaining shoulder stability.
.)~ Elevation of the arm involves motion at both
the glenohumeral joint and the scaplliothoracic articulation.
?;S) Movements of the spine assist the shoulder in
positioning the UppCt' extrcmity in space.
{g) The muscles around the shouldcr contribute lO
st~~'bility via a barricr effect by producing compressive forces on the glenohumeral joint and byeccentric contraction.
~Q';The glenohumeral joint is a major load-bearing
joint with forces equivalent to one-half body weight
produced when holding the arm in an outstretched
position.
REFERENCES
:\gur, A.M.R .. lee. \1., &. Grant. J.e. (1991). :lIla ... o":lwllomy
(9th NI.). Bahimorl..': Williams &. Wilkins.
Baml..'s, D.A. &. Tullos, H.S. (1978). An analysis of 100 symptom'llie b<tsl..'ball playt.'fs...\11/ J Sports Mel!. 6. 62-67.
Boonc. D.C. &. :\zcn. S.P. (1979). Nornwl range of Illotion of
joints in male subjl.'cts. 1 Bow: loiHt S/frg. 61. 756-759.
Colaehis, S.C. Jr. &. Strohm, B.R. (1971). Effl..'ct of suprascapul;lr and axillary nervc blocks on muscle farer: in upper
I..'.xtrl..'milv. Arch Phvs :\Jed RciUlbil. .12, 21-29.
Cooper, D.E., O'Briel~, 5.1., Amol:Zk:". S.P., 1..'1 ;;d. (1993). The
structurc and function of the coraCOllllllll.'ral ligamcnt: An
anatomic and mil.:roscopic study. J Shoulder Elboll· SlIrg. 2,
70-77.
Curl. L.A. &. Warrcn, R.F. (J 996). Glenohumeral joinl stabil·
ity: ScI",'ctin' culting studies on the slatic: c<\psubr re·
siminis. Clill Onhop, 330, 54-65.
Delee J. & Ora, D. (199.:1). Onhopac:l!ic Spons Mc:dicillt·.
Prillciplt:s awl 1)raelicc (p. 464). Phibdelphin: W.B.
Saunders.
Dc P.dlll':\. A.F. (1983). BiollH.'chanics of the shoulder. In
SI/rgery
tlte Shollider (3rd cd .. pp.65-85). Philacldphia:
J,B. Lippincolt.
or
Fukllda, K.. Cr;lig. E.\·.. All. K.. t·l ;d. 11{J~6i. Bihmt·
:-llId,- e,l Illl· ligallll·lll('ll:' :-~:-tt·1ll (If thl' :lcnHnicH
jllill!..1 HfJlil' .If/ill; Sltr::. MS. ·t3.t-~.tO.
Gihh. T.IL Sidle,. J.:\ .. lbn·~lll;tll. D.T.. d :11. (1991)
ft...:! (d- l"~lp:-lIlar \l·lIling oil gk·lIohllllkr:d 1:,'\iiY.
!/f(ip. 268. 110-11i.
l!ollillShl·"d. \\'.11. (IY6t)). ·\lIli//llllf /n,. Sllr.::OJl/\ (VoL
York: Harp..·.. ...\: Ih,\\.
lIe,\,-dl. 5.:\1.. IIlHdlcr:-kg. :\ ..\\.. S..· ga, D.lI .. l:I :11
Cbrifil:;ltion 01 till.· nd .., Io! Ih~· :-upr:l:-pill:IlU:- n
:-houldl·r 11llKliclll. 1 H"'I/, 11'iHI Slfl~. 00:\. 3YS-·W
[nm:\ll. V.T.. S;\Illllk·r:-. J.I3 .. & :\blwll. L.C. (19.t ..+.!.
lifl!\S nn Ihe llln;.:lillll III till· :-houldn jlllnl . .1 !J
SllJ·.~.
26..\. 1-30.
Iloi. E., I-bll. I-I.e.. l\: All. h.\. (199{\1. BiOlllt'ch:lnica
g;llioll of lhl: gkllohllllll:r;d join!.) SltlJ//lcla Uhow
40i- ..L~~.
Iloi. E.. \lol/.kin. ~.E . .\IMr..·.\. B.F.. d :tl. {IY9-ii. SI
illlKlioll Ilf tht· long hC:ld of Iht· hic..·p:- ill till· han
po:-,itilln. J Sltould..r FlblOw Sur;.!.. 3, 135-1-I:!.
hoi. E. .\l(llzkin. \:.E. . .\1{)lTt·'. 13.17.. l'l ~tl. (1992).
inclill:ltiol\ ~ll\d illk·ri(lr :-1;lbilil, 1'\ lilt· ... huuldCl".
tit'/" r:1l}('1~' Sun.;. I. 131-1.,9.
"-l1m;lr. \'.1'. & lbb:-uhr;ltll;tlli:llll, P. (1')S51. Tilt· rC
n\I,:,ph~·rjt· pr~· . . :-t1r~' ill :-1:lhili/ing lllv :-h(,tlllkr: i\
ll\~·llt:d ~tltd\. J nUll<' Jpiur Sill.!.:. (,iii. 71'L·711.
L'llilll'11HI. t'. i 1'J~7). Kin~·~i()I(I~'.\' (Ii I h\.· sil(llillkr jo
K(,ih~·1.
~·t a!. i. E.d~.I. ,I,,;/inlildl'l' N'Pi!;Cl'iJl,'IIi.
Sprinl,'n· \"nl:q!. I 'J~ 7.
Lippitt. S,B .. \·"JHkrho()ft. .J.F.
GICI1(lhllI11~'!":\1 :;t~lhijit'
q\l'lnlil'lli\~· :ll1;I!.,:;i:, .
.1
J[;\I'1'i~,
S.L.
~·t ~i
('(JJICI\'il.\·l:olllpr,,
SJIO/lldrJ' IJ;'II\\' Sur::, }, 2
I"rllill
[.tll·,t:'. I).B, (1973). Biolll(·Ch:ltli\.'s 01 tht· :;hntlltlt·r in
Sw·:,;. 107. ~2~ .
.\bbC:l. E. Fl!. F.. \.\: 1[;1\\ kill:-. R. 1Fd ..... \ \ 19921. rh,·
:1 J3a!IiIiC" iI! _\//Ohilizy rll/{; Stahility. RO:-Clllill1l. I
I\';II!. Colorado: \\'lIl"bhop Suppllrkd b., dl~·
:\(adt'my (If Orth(Jp~lt·dic Sllr~"·{)I\:-. Ih..• X<llional
of r\nhrili:- alld .\llI:-nllo~k~·k!:t1 Skill I)i~t·a:-;..,:-. t
i t-:Ill ShouldtT ;1Ilt! ElbO\\ Suq:!t·OII:-. Ih ..· On!lopa
:-t·~lr\.>h and E.hIGII il'" F('lllltbi iOll.]
.\lilffcy. B.F. & :\n. I\..\:. i 1990), Biolllt·dl:lllk:o, (If the
III c.:\. ROi..·k\\(I(,t! \'-. F.:\ . .\1~lt:-t·1l III {Elk). The
lpp. 10~-2.t:;L rhibddphi~l: \\".B. S;illl1(lt-r:-.
.\I(lor..·. h.L. (19t)\ji. Cliuh:{:lly O,-it'lth'd ..Ii/ili/IJllY 1
l)hibtklphi"l: LippilKllti \\'illl,tlll:- & \\"ilkin:-.
.\hUTt·'·. B.F. l\: :\11. "-.X. (1990). BiCll11t'I.:h:lllk:-. of l
tit-I'. In C:\. Ritt"kW(IOd l\: 17.:\. '\\;il:-t·1l 111 {Ed~.). rI
d,·,: PhiliHk·lphia: W.B. S:llln.kr....
.\lurr:I\·. ;\-I.P.. GMt.. D.R .. Cardlll·r. C. .\I.. ('1 ,d. l19~5
tin' motion ;llld Illll:-r.:k :-lrt·ngth of normal
\\Olllt'll in Iwo ;l~e ~rollp:-,. Clill Onhflp. /92, 193.
:\cl..'r. C. (1990i. Shoulder UC·(·flll:'In/oiIlJl (p. ]9).
phi;t: W.B. Saundns Co.
O'Brit·n. Sol .. \ne:- . .\1.C. "l1Hli.:l.k~. 5.1'.. d al. (1
:lnalomy and hi:'lology 01 Ih ..· inkriof gknohullw
lIh:nl cOlllpkx of I lit· shoulder. :\111 J Sports
579-3$4.
Pagn:llii. ,\-l.J .. Dt'llg, X.Il .. Wan\'Il. R.F.. l"l ;11. (1995),
k:-;iolls of the superior 1l<ll"Ii(lli of lilt' glt'noid b
glenohullleral tr'lmlali(JIl:- . .I1J<"h,·.Il,1illl SIlI".:;. 77:\.
Popp(,ll, N.K. & Walh'r. P.S. (1978). rorcl.'s al Ih(' giL'nohllnlt:r,1! join! in ,lbdllClioll. Clill OnJ/lJp. 135. 165-[ 70.
poppell, N.K. & W~lkl.'r, IJ.S. (1976). NOl"m~1 .111<1 abnormal
motion of thl.' shoulder. J BOlle Joi/lt SlI(!!.. 58..\, 195-20 I.
Rochmod. c..-\. h. (1975). Disloc:.uions about lht: shoulder.
In C.A. Rockwood Jr. ..\: D.I J _ Grl.·cn (Eds.), Fmc/m't's (1st
,'j l.'d .. Vol I, PI'. 624-815). Philadelphia: J.B. Lippincott.
C. ..\: Malst:ll. F. (1990). The ')·!lo/ddt·" (p. 219J.
Philaddphia: W.O. Saunders Co.
S.R. (Ed). (1994). Orthopaedic Basic .'idCllce (p. 327).
Rosemont, IL: AAOS.
L.J., Flaww, E.L., Bigliani. L.U., C[ HI. (1992). Articular geonll:lr~'
IhL' glenohumeral joillt. Clill Ol"lhop,
or
28;, 181-190.
Sh.:inbt:ek, J., Lilje:lIq\'isl. U., 6: Jcrosh, J. (199-0_ The
contribution
to
<In[aior shoulder stnbility. J Sho/!lder
bOIl'SlIrg, 7(2J.122-6.
Tibont:. J .. P;Hd, R., Jobe:, F.W.. t:l a!. (199-0. The Shoul
fUIKtional analOmy. hi(1/1h:<:hanics and kinesiology. In
De:Lt:1: 6: D. Dra (Eds.), OrillOpa.:dic SPOrtS ,\Iedic
IJhilnddphia: W.B. Sallnde:l"s Co.
W:ll'I1C.... J.P, (1993). Thl; gross all~lIomy of the joint surfa
lig<tll1<.'nts, labrum :lIld c.::lpsuk. In F.A. Matsen, F.t!. F
R.J. H.l\\·kins (Eels.), Thl.' Should{'/": A !3alaJ/n: of .\fob
and Flllictioll (PI'. 7-17). Rosemont, It: AAOS.
Wnrncr, J.P., Dcng, X.I-! .. \Varren, R.E et al. (1992). SU
l::lflsllloligamenlOlls rt.'siraints to superior-inferior tr:lf1
lions of the g!t:!lohulllcral joint. Alii } SPOI'1S ,\Jed,
675-678.
an.atolll)' or Ihl: gle:nohlimcrallig~lIne:nlOlis compkx and its
:/;'
~
L
•
O>-~
'"
"'"!'.,-_-.!'.!','!'.A__"'zg!,!,,,.,,!,!,,"""'_.","_!,,,~".B'" "'_........"_...,;"t.~".,,__,,,",_"_,·c"-, ___·,"'.", "_--",,---~. ·"",~c_,,,:-,-,,,,~,~,,_,,,,;;o..,,,,,,~,?,,_,,~,,,_--- __- __,,,,,,"7.2'}"'...!,,",,J.7!!J~'!;?3?WC'--"; _ ..!_".xa"'..", ;. ""'.."'""_,
Biomechanic
of the Elbow
Laith M. Jazrawl~ Andrew 5. Rokito, Maureen Gallagher Bird
Joseph D. Zuckerm
Introduction
Anatomy
Kinematics
Carrying Angle
Elbow Stability
Kinetics
Electromyography
Elbow Joint Forces
Artieular Surface Forces
Calculation of Joint Reaction Forces at the Elbow
Summary
References
. ,
•
f ,,Introduction
~,
IX
~/
t
Wi,
t,
!
·i
The elbow is a complex joint that functions as a fulcrUIll for the forearm lever system that is responsi-
~•. ",.'
~
j
I
I
. hie for positioning the hand in space. A detailed lin-Hi;:Xderstnnding of the biomechanics of elbow function
J,;':
\
I
,is essential for the clinician to e1Tcctivcly treat
'pathological conditions affecting the elbow joint.
~"'"
~:'.•,.,· .;;'.A
/'l ato III V
Wi
I't: The
elbow joint complex allows
types of molion: nexion-extension and pronation-supination.
~
,•.... '.'
.,/
1\\10
i,
The humcroulnar and hUlllcroradial articulations
~
allow elbow Oexion and extension and arc classified
i
as ginglymoid or hinged joints. The proximal ra-
~,:
dioulnar articulation allows forearm pronation and
i
supination and is classified as a u'ochoid joint. The
elbow joint complex. when considered in its en~
tircty. is therefore a tl'ochleoginglymoid joint. The
trochlea and capitellum of the distal hUl11cnls arc
internally rotated 3 to 8' (Fig. 13-IC) and in 94 to 98'
of valgus with respect to the longitudinal axis of the
humerus (Fig, 13-IA). The distal humerus is anteriorly angulated 30° along the long axis of the
humenls (Fig. 13-1 B). The a.-licular surface of the
ulna is oriented in approximately 4 to 7° of valgus
angulation with respect to the longitudinal axis of
its shaft (Fig. 13-2;\).
The distal humerus is divided into medial and lateml columns that terlllinate distally with the
trochlea connecting the two columns (Fig. 13-3). The
medial column diverges [Tom the humeral shaft at a
45° angle and ends approximatel,Y 1 cm proximal to
the distal end of the trochlea. The dislal one third of
the medial column is composed of cancellous bone,
is ovoid in shape, and represents the medial epicondyle. The lateral column of the distal humerus
diverges at a 20° angle from the humerus at the same
level as the medial column and ends \'.lith the capitelIUlll. The trochlea is in the shape of a spool and is
comprised of medial and lateral lips with an intervening sulcus. This sulcus articulates with the semilunar notch of the proximal ulna. The articular surface of the trochlea is covered by hyaline cartilage in
an arc of 330°. The capitellum, comprising al0105t a
perfect hemisphere. is covered by hyaline cartilage
forming an arc of approximately 180°.
Tbe mticular surface of the ulna is rotated 30' poste·
Jiorly with respect to ilS long ax.is. This matches the 30°
anteIior angulation of the distal humerus. which helps
pl"Ovide stability to the elbow joint in full extension
I
I
rf
i,
i
~i
~.
K
~~
t'
~
~
i
.,.
,
~.
~
I
::
$
ii;
(
~..
<
I.
l
~
I
~
r:
t
~
I
t
r
30'
B
A
c
Angular orientation of the distal humerus in the anteroposterior (A), lateral (B). and axial (C) projections.
4'
~.
I
A
B
Angular orientation of the proximal ulna in the anteroposterior (A) and later