STATISTICAL COMPARISON OF CUSTOMIZED AND TRADITIONAL KNEE IMPLANTS
USING FINITE ELEMENT ANALYSES
A Thesis by
Atta-Ur-Rehman Hashmi
Bachelors of Science in Mechanical Engineering, Wichita State University, 2001
Submitted to the Department of Industrial and Manufacturing Engineering
and the faculty of the Graduate School of
Wichita State University
in partial fulfillment of
the requirements for the degree of
Master of Science
July 2007
© Copyright 2007 by Attaurrehman Hashmi
All Rights Reserved
COMPARISON OF CUSTOMIZED KNEE IMPLANT WITH TRADITIONAL KNEE
IMPLANT USING DIFFERENT ORHTOPAEDIC MATERIALS
I have examined the final copy of this thesis for form and content, and recommend that it be
accepted in partial fulfillment of the requirement for the degree of Master of Science with a
major in Industrial Engineering.
_____________________________________
Gamal Weheba, Committee Chair
We have read this thesis and recommend
Its acceptance:
_____________________________________
S. Hossein Cheraghi, Committee Member
_____________________________________
Charles Yang, Committee Member
iii
TABLE OF CONTENTS
Chapter
Page
1.
INTRODUCTION…………………………………………………………………….
1
2.
LITERATURE REVIEW……………………………………………………………..
3
Materials in Orthopedic Surgery……………………………………………....
2.1.1 History………………………………………………………………....
2.1.2 Material Properties…………………………………………………….
2.1.3 Physical Properties…………………………………………………….
2.1.4 Biological Properties…………………………………………………..
2.1.5 Metallic Orthopaedic Materials.……………………………………….
2.1.6 Stainless Steel………………………………………………………….
2.1.7 Cobalt Chrome Alloys…………………………………………………
2.1.8 Titanium and Titanium Based Alloys………..………………………...
2.1.9 New Materials for Knee Implants……………………………………..
The Knee……………………………………………………………………....
2.2.1 Anatomy……………………………………………………………….
2.2.2 Knee Problems………………………………………………………...
2.2.3 Knee Diseases…………………………………………………………
2.2.4 Knee Injuries…………………………………………………………..
2.2.5 Knee Problem Treatment……………………………………………....
Knee Replacement…………………………………………………………….
2.3.1 Brief History…………………………………………………………...
2.3.2 Knee Implants………………………………………………………….
Customized Knee Implant Design…………………………………………….
3
3
5
6
10
12
13
14
15
17
19
19
26
27
29
33
34
35
36
42
COMPARISON OF KNEE DESIGNS ….………………………………...................
47
2.1
2.2
2.3
2.4
3.
3.1
3.2
3.3
3.4
3.5
3.6
Research Methodology………………….…………………………….………
Customized Femoral Component Model....…………………………….……..
Traditional Femoral Component Model....…………………………………....
The General Factorial Design…………………………………………………
Finite Element Analyses.…...…………………………………………………
Statistical Analysis.……………...…………………………………………….
47
48
53
55
56
62
DUSCUSSION & CONCLUSIONS…..……………………………………………...
66
REFERENCES……………………………………………………………………………..
69
4.
iv
TABLE OF CONTENTS (continued)
Chapter
Page
APPENDICES………………………………………………………………………..........
A Stress Level max Comparison Charts……………...………………………….
B FEA of Traditional Knee Design (Stainless Steel)……….……….…………..
C FEA of Traditional Knee Design (Cobalt Chrome)...………..………………..
D FEA of Traditional Knee Design (Ti6Al4V)…….……………………………
E FEA of Traditional Knee Design (Oxidized Zirconium)...…………………....
F FEA of Customized Knee Design (Stainless Steel)…………….……………..
G FEA of Customized Knee Design (Cobalt Chrome)...……….………………..
H FEA of Customized Knee Design (Ti6Al4V)……..…………………………..
I FEA of Customized Knee Design (Oxidized Zirconium)...………….………..
J ANOVA Residual Plots……………………………………………………….
75
76
77
80
83
86
89
92
95
98
101
v
LIST OF TABLES
Table
Page
2.1
Mechanical properties of Metal Alloys and Bone…………………………………. 13
2.2
Mechanical Characteristics of Orthopedic Alloys…………………………………
17
3.1
Design Factors and Their Levels…………………………………………………..
56
3.2
Load Cases…………………………………………………………………………
58
3.3
Material Properties…………………………………………………………………
59
3.4
Design Matrix with FEA Results (Measured Response)…..………………………
60
3.5
ANOVA Table………………………………………….………………………….
63
vi
LIST OF FIGURES
Figure
Page
2.2.1
Bones of a Knee…………………………………………………………………...
20
2.2.2
Knee Components…………………………………………………………………
21
2.2.3
Articular Cartilage………………………………………………………………...
22
2.2.4
Ligaments in the knee……………………………………………………………..
22
2.2.5
Movement controlled by ligaments……………………………………………….
23
2.2.6
Menisci of the knee………………………………………………………………..
24
2.2.7
Knee Tendons……………………………………………………………………... 25
2.2.8
Bursae of Knee Joint………………………………………………………………
2.2.9
Walk mechanism………………………………………………………………….. 26
2.2.10
Osteoarthritis Knee Joint………………………………………………………….. 27
2.2.11
Rheumatoid Arthritis Knee Joint………………………………………………….
29
2.2.12
Chondromalacia Patella…………………………………………………………...
30
2.2.13
Meniscus Tear……………………………………………………………………..
31
2.2.14
ACL & PCL ligament injuries…………………………………………………….
31
2.2.15
Medial & Lateral collateral ligament injuries……………………………………..
32
2.2.16
Tendinitis………………………………………………………………………….. 33
2.3.1
Components of a knee implant……………………………………………………. 37
2.3.2
Fixed Bearing Knee Design……………………………………………………….
2.3.3
Mobile Bearing Knee Design …………………………………………………….. 39
2.3.4
Unicompartmental Knee Design ...……………………………………………......
vii
25
38
40
LIST OF FIGURES (continued)
Figure
Page
2.3.5
PCL-Retaining Knee Design……………………………………………………… 40
2.3.6
PCL-Substituting Knee Design……………………………………………………
41
2.4.1
CT-images of author’s knee……………………………………………………….
43
2.4.2
3D model of distal femur………………………………………………………….
44
2.4.3
Iges-curves imported into Pro-E…………………………………………………..
45
2.4.4
Spline curves created in Pro-E…………………………………………………….
45
2.4.5
Customized Femoral Design………………………………………………………
46
3.1
Research Procedure………………………………...……………………………... 48
3.2
Imported Knee File in Catia (different view angles)……………………………...
3.3
Sliced cross-sections of femoral component……………………………………… 50
3.4
Solid Cross Section of Customized Knee Design Model…………………………. 51
3.5
Femur bone attached to the customized knee implant…………………………….
52
3.6
Industry Design Standard for Femoral Component……………………………….
53
3.7
Parametric Description of Femoral Component Geometry……………………….
54
3.8
Solid Cross Section of Traditional Knee Design Model………………………….. 54
3.9
Femur bone attached to the traditional knee implant ……………………………..
3.10
Loads and Boundary Conditions at 0° Angle……………………………………... 57
3.11
Loads and Boundary Conditions at 45° Angle……………………………………. 57
3.12
Sample FEA results for Traditional and Customized Implants…...………………
61
3.13
Design – Angle Interaction………………………………………………………..
64
3.14
Weight Effect Plot………………………………………………………………… 65
viii
49
55
ABSTRACT
According to the National Center for Health Statistics in 2004 over 478,000 people
underwent Total Knee Replacement surgeries in 2004. The number is rapidly growing due to
advances in implant surgeries, knee implants, and longevity but is limited to patients older than
60 years. Knee implant surgery failures have also increased which has lead to increased number
of revisions. Recent attempts have been made to design and optimize customized knee implants
for young patients and provide a solution for current failures in knee replacements.
The objective of this research is to compare the femoral components of customized and
traditional implant designs using femur bone interface. The comparison was made using three
load levels of body weight acting on the femur bone with attached femoral components at two
gait angles with four orthopaedic materials. Statistical analysis was conducted using a general
factorial design to quantify the effect of these factors. Finite Element Analysis (FEA) was
performed to measure the maximum stress level (
max)
for design comparison. The statistical
analysis concluded that the two-factor interaction involving knee design and angle had a
significant effect on the average stress level. Based on the assumptions the FEA results indicated
that the femoral component of the traditional knee implant design was better than the customized
knee implant design with respect to maximum stress level (
ix
max).
CHAPTER 1
INTRODUCTION
The number of knee replacement surgeries are rapidly growing due to increased life span,
success of many implant surgeries, more human activities, among others. On the other hand the
problems related to these implants have also increased. The factors contributing to these
problems are loosening of components, mechanical failures of the implant or failure of the bone
structure. Traditional implant procedures involve using “Standard” implants. These knee
implants are mass-produced in limited sizes and designs. It is up to the surgeon to choose the
best fitting implant for the surgery. Surgeons adopt the patient bones to the standard by removing
valuable naturally grown bone structure. The main problem with traditional knee implants is the
limited life span causing patients to go through several revisions during their lifetime. The 2006
Annual Report published by the Swedish Knee Arthroplasty Register indicates that most
revisions are due to the loosening of the implant components which is mainly caused by poor
initial fit between the bone and the implant [1]. The average age of these patients is above 68
years for both male and female. Major factors responsible for these revisions included polythene
wear, loosening, and instability.
Recent developments in fields of Rapid prototyping (RP), Reverse Engineering (RE) and
Image Processing (IP), have evolved in a new field of Medical Applications of Rapid
Prototyping (MARP). RP is the process of converting CAD files into physical models “rapidly”.
Rapid Prototyping technologies were originally developed for the fast production of physical
products based on Computer Aided Design (CAD) models. This technology is widely used in
product design of some industries such as aircraft, automobile, and tools. Many natural objects
1
such as bones and tissues are not accessible to direct inspection and handling but are of great
scientific interest. Combining an imaging technology like computer tomography (CT) with a
rapid prototyping (RP) technology like stereolithography can generate 3D hard copies of these
objects for use in such applications as surgical rehearsal and planning, communication between
medical staff and the patient, custom implant design and casting.
Harrysson [2] proposed a proof-of-concept to customize and optimize knee implants for
young patients by combining technologies like CT, reverse imaging, rapid prototyping and
robotic surgery. Patient’s knee CT-data was converted to a CAD model using imaging
processing and CAD software. The CAD model was used to design, customize and optimize the
femoral and tibia components of the knee implant. The research provided a complete
methodology to develop customized knee implants.
The objective of this research is to compare the femoral component of the customized
and traditional knee implant designs with bone interface using four different Orthopaedic
materials: Stainless Steel, Cobalt Chrome, Titanium (Ti6Al4V) and Oxidized Zirconium. The
customized femoral component, supplied by Harrysson [2], was analyzed and compared with
traditional femoral component attached to the femur bone using Finite Element Methods. The
analyses were performed using three load levels at two gait angles and four Orthopaedic
materials. Statistical analysis was performed using a general factorial design to study the result
(Maximum Stress level -
max)
from FEA and to provide conclusions on the study. The next
chapter (Chapter 2) provides a literature review for Orthopaedic materials, knee, knee
replacement, and the methodology behind customized knee implants. Chapter 3 discusses the
research behind design comparison, FEA and statistical analysis. Discussion and conclusions are
outlined in Chapter 4.
2
CHAPTER 2
LITERATURE REVIEW
2.1 Materials in orthopaedic surgery
Metal alloys have been the materials of choice since the beginning of orthopaedic
surgery. Orthopaedic materials must satisfy the mechanical, biological and physical requirements
of their intended use. This section provides the literature review for orthopaedic materials
involved in implant surgery.
2.1.1 History
One of the first internal fracture fixations occurred in 1775. An iron wire was used to
support the bone. It was surrounded by controversy due to the reaction of the tissues around the
implanted wire. At the time, the reaction was solely attributed to the implant [3].
In 1866, Hansmann became the first surgeon to use metal plates and screws to repair
bone fractures for internal fixation. By the end of the century carbon steel nails coated with gold
were being used for surgery [4]. In 1892 Levert used lead, gold, silver and platinum wires in
dogs and found that these metals did not have the desired mechanical properties. He also
concluded that without anesthesia, human patients could not endure long surgeries in order to
implant meaningful prosthesis or fixation devices [5]. Lambotte in 1909 compared several metals
and found gold or nickel plated steel had the most satisfactory in mechanical and corrosion
resistance qualities [6].
Vanadium steel was the first metal alloy developed by Sherman specifically for human
use as fracture fixation plates. He reported the use of vanadium steel plates and screws in 1912.
Early on the attempts to employ surgical implants were hampered due to the limitation of
3
available materials [7]. Metallic cobalt started to find some industrial use at the beginning of the
century. Since it’s pure form is not particularly ductile or corrosion resistant, a series of cobaltchromium and cobalt-chromium-tungsten alloys were developed between 1907 and 1913 which
promised good corrosion resistance [8].
Further research continued until 1924 when A. Zierold published a study on tissue
reactions of various metals implanted in dogs to determine their comparative biocompatibility.
His findings pointed out that satellite, a cobalt-chromium alloy, was highly biocompatible. Iron
and steel were found to corrode rapidly leading to resorption of adjacent bone. Copper,
magnesium, aluminum alloy, zinc, and nickel discolored the surrounding tissue while gold,
silver, lead and aluminum were tolerated but inadequate mechanically [9].
From 1923 to 1938 Dr Smith-Peterson performed preliminary trials using several
materials including glass (1923), viscoloid (1925), pyrex (1933) and bakelite (1938) as implant
materials. These materials did not work due to poor mechanical properties [10]. In 1926 the first
stainless steel used for a human implant was 18% Chromium-8% Nickel (type 302 in modern
classification). It was noted to be much stronger and highly resistant to corrosion than vanadium
steel. Later, the 18Cr-8Ni was supplemented with a small amount of molybdenum and silicon to
improve the corrosion resistance in salt water. This became the standard metal alloy for medical
implants as type 316 Stainless Steel [4, 11].
During the early 1930s an alloy called Vitallium with a composition of 30% chromium,
7% tungsten and 0.5% carbon in cobalt was employed for the preparation of metallic dental
castings [8]. Alvin Strock placed the first successful oral implants in 1937 at Harvard University.
Strock published a paper on the physiological effects of cobalt-chromium-molybdenum alloy
4
(vitallium) in bone, and thus placed a series of Vitallium implants into test animals and humans
[12, 13].
Venable in 1938 reported the first use of Cobalt-chromium alloys. Since then they have
been developed and proven effective as corrosion-resistant orthopaedic surgical implants [5]. By
the 1940s stainless steel was well established as the choice of material for medical implants. In
the 1950s, a low carbon version of 316 (type 316L) was developed to reduce the carbon content
from 0.08 wt% to 0.03 wt% in order to improve corrosion resistance further [14].
Later on a new alloy is introduced into orthopaedic surgery. Titanium and its alloys had
shown excellent inertness in the human environment making it a corrosion resistant implant in
the human body. This capability was confirmed after surgical implants (screws and plates) were
tested on human subjects in the early 1951 by Leventhal. Subsequently a variety of titanium
orthopaedic implants and devices were introduced [15, 16]. Research is continuously ongoing to
find improved alloys with preferable properties.
2.1.2 Material Properties
Orthopaedic implants are more advanced today when compared to their predecessors.
The materials (usually known as biomaterials) are highly developed. Examples of biomaterials
include Titanium, Cobalt-Chrome, Stainless-Steel, and polyethylene. In the United States the
FDA requires extensive testing before a new material may be used in an orthopaedic implant.
The materials most commonly used have a long history of clinical use with great success [17].
There are many different biomaterials, but there is no single material that is best for all implants
and all patients. The specific requirements of an implant material vary depending on how the
implant is designed to be used. Also, like medicine, biomaterials can produce side effects like
microscopic debris, increased ion levels in the blood or urine, or inflammation. For these
5
reasons, doctors evaluate the patients individually and carefully consider the material used to
manufacture the implant, along with its design.
During daily activities an implant may encounter mechanical forces that tend to push,
pull, bend, scrape, or cause its parts to rub together. These forces can cause the implant to break
or wear out over time. The mechanical properties of a biomaterial can best be described by its
modulus of elasticity, ultimate tensile strength, elongation to failure, and fracture toughness [18].
− Modulus of elasticity describes the stiffness of the material and is usually obtained from the
slope of a stress-strain diagram.
− Ultimate tensile strength describes the ability of the material to withstand a load before it
fails.
− Elongation to failure describes how much strain the material can bear before it fails.
− Fracture toughness is an important measurement of the material’s resistance to crack
propagation.
The materials are also subjected to the many natural chemicals inside the body. Although
normal, some of these chemicals may tend to corrode some materials. In order for an implant to
perform under these conditions, it must be made from materials that can withstand these forces
and chemicals. Whether an implant is designed to replace a joint, or help repair a fracture,
several physical and biological characteristics are important when selecting the material for the
implant.
2.1.3 Physical Properties
The ideal implant material should have physical characteristics that match those of the
bone it is replacing or reinforcing. Since orthopaedic implants are attached to human bones, they
6
must work with the patient’s bones to restore function. This usually requires a balance of
physical characteristics. For example, human bones are strong but flexible. This combination
helps them withstand forces as high as several times the human weight without breaking [19].
Some of the physical characteristics for medical implants are explained below:
Fatigue Failure
Fatigue failure is probably the primary concern for orthopaedic applications of surgical
implants. Biomaterials used in the design of medical devices are subjected to high stresses and
high cycle loading. These demanding loading conditions results in failure of metals, ceramics,
and polymers. The fatigue process depends on stress rather than load, therefore partly explaining
the success of congruent total hip replacements and thick polymeric bearing components which
allows for a larger contact area [18, 19].
Fatigue failure in a material can and often occurs under loading conditions where the
cyclic or fluctuating stress is below the tensile strength. The resulting stress which is below the
ultimate tensile stress and in some materials even below the elastic limit causes failure. Fatigue
failure occurs at significantly lower life as the applied stress increases. For some metals or alloys
such as steels, the stress level is observed below which failure does not occur. In most steels this
“fatigue limit” is empirically observed to be about 0.5 of the ultimate tensile strength. Therefore,
by selecting a high-strength material the allowable working stresses may be increased [21].
In metals the crack usually initiates at the surface and grows slowly during the functional
life of the specimen. This crack may propagate in a zigzag fashion. Each stress cycle further
accelerates the crack growth. Eventually when the residual section is sufficiently small and
highly stressed it fails by conventional overload mechanisms. A number of factors effect fatigue
life. Surface notches or holes severely affect fatigue strength. Defects such as scratches or
7
corrosion can reduce the strength by up to 20% [22]. Fatigue failure of surgical alloys is
gradually becoming the principle cause of implant failure. Fatigue failure is usually associated
with poor design, workmanship or handling. Attempts have been made to eliminate stress
concentration such as crevices, corners or other irregularities. In practice however, even if the
implants arrives in the theatre in pristine condition, scratches will inevitably occur during
surgical implantation. The best hope for improvement rests on the possible introduction of a
variety of fiber reinforced composite materials that show remarkable resistance to fatigue.
Wear
Materials used in the fabrication of implants are subjected to wear. Wear of materials and
devices has been shown to be detrimental to their long term success resulting in implant retrieval
and revision. One of the most dramatic impacts of the wear of materials and its consequences is
observed with artificial joints. Wear has emerged as a central problem limiting the long-term
longevity of total joint replacements [23]. Ultra-high-molecular-weight polyethylene
(UHMWPE) wear debris has been shown to trigger an osteolytic reaction which leads to implant
loosening [24]. Wear is a process resulting in the progressive loss of material involving many
diverse mechanisms and phenomena which are often unpredictable. The wear process of
materials is predominantly governed by their mechanical and/or chemical behavior. More often
than not, the wear processes like adhesive wear, abrasive wear, and corrosive wear do not act
independently. However, even though several wear mechanisms are involved, it is often the case
that one particular mechanism dominates.
Unfortunately surface wear of an implant results from its use, and therefore, cannot be
avoided or eliminated. Because wear is a limiting factor in the successful outcome and lifetime
of an implant, it is of the utmost importance to characterize the wear resistance of materials used
8
in implant design, and the effect of the design on wear [19]. The wear and frictional properties of
materials are dependent on tribological conditions of the implant. Their investigation involves
many parameters such as wear rate, wear mechanisms, transition between initial and steady-state
wear, and generation and geometry of wear debris. The physical and mechanical properties of the
materials, the environmental and operating conditions, and the geometry of the wearing bodies
are determining factors for these parameters.
Historically, the evaluation of materials has resulted in the development of new materials
or the use of surface treatments to improve their frictional and wear properties. Surface
treatments are applied on the implant materials to enhance the hardness and wear resistance.
Another approach to improve the wear performance of total joint replacements is through
lubrication optimization [20].
Corrosion
Corrosion is one of the major processes that cause problems when metals and alloys are
used as implants in the body. Wear and release of soluble products from implant materials results
from the degradation of materials. In the case of alloys, soluble ions and compounds are released
due to corrosion. Corrosion is defined as the unwanted chemical reaction of a metal with its
environment, resulting in its continued degradation to oxides, hydroxide or other compounds.
Metallic surfaces in contact with body’s fluids corrode. Their surface dissolves and the dissolved
metals enter the circulation. The concentration of the metals (Cobalt, Chromium, or Titanium) in
the blood increases. To a certain extent metal ion release from alloys in the body always occurs,
although the ions released are relatively negligible and generally this low level of contamination
is not classed as corrosion [25].
9
Corrosion can have two effects. First the implant may weaken and the premature failure
will result. The second effect is the tissue reaction leading to the release of corrosion products
from the implant. No metallic material is totally resistant to corrosion or ionization within living
tissues. In vivo studies have shown that the implantation of the devices of most alloys
significantly increases the concentrations of various ions adjacent to the tissues. Moreover, once
a foreign material is implanted, there are several ways in which the body may react unfavorably.
The presence of the implant may inhibit the defense mechanisms of the body leading to
infection, necessitating the removal of the implant [26]. If infection does not occur or is
controlled, the tissue response may range from mild edema to chronic inflammation and
alteration in bone and tissue structures. This necessitates that the materials used in making
implants must be inert or well tolerated by the body environment [27].
Most of the traditional methods of controlling corrosion cannot be used for surgical
implants as the environment within the human body is fixed. The only methods available are to
fabricate the implants from a corrosion-resistant alloy or to use a coating - either of which must
be able to withstand any abrasion and wear to which the device may be subjected [28].
2.1.4 Biological Properties
Along with physical characteristics, the biological properties are just as important for
orthopaedic materials. Biological characteristics refer to the biological effect the material has on
the body, as well as the effect the body has on the material. The human body is a harsh
environment for metals and alloys having to be in an oxygenated saline solution with salt content
of about 0.9% at pH ~7.4, and temperature of 37±1°C (98.4°F) [29]. All the surgically
implantable metallic materials, including the most corrosion-resistant materials, undergo
10
chemical or electrochemical dissolution at some finite rate, due to the complex and corrosive
environment of the human body.
Biocompatibility
Biocompatibility implies the ability of the material to perform effectively with an
appropriate host response for the desired application. The host response to biomaterials is
extremely varied, involves a range of different mechanisms and is controlled by factors that
involve characteristics of host, material and surgical procedure. These responses themselves
constitute a significant component of the phenomenon of biocompatibility [30].
The biocompatibility of an artificial material in the body is extremely complicated,
involving processes traditionally belonging to medical science, surface science, materials
science, and molecular biotechnology. There are various components that are involved in the
biocompatibility process. Biocompatibility refers to the totality of the interfacial reactions
between biomaterials and tissues and to their consequences [31].
Biological evaluation of medical devices is performed to determine the potential toxicity
resulting from contact of the component materials of the device with the body. The device
materials should not, either directly or through the release of their material constituents: (i)
produce adverse local or systemic effects; (ii) be carcinogenic; or, (iii) produce adverse
reproductive and developmental effects. Therefore, evaluation of any new device intended for
human use requires data from systematic testing to ensure that the benefits provided by the final
product will exceed any potential risks produced by device materials [32]. FDA recognizes the
standard ISO 10993 for biological evaluation of medical devices. This standard provides
guidance for selecting the tests to evaluate the biological response to medical devices. When
selecting the appropriate tests for biological evaluation of a medical device, the chemical
11
characteristics of device materials and the nature, degree, frequency and duration of its exposure
to the body must be considered.
2.1.5 Metallic Orthopaedic Materials
The main requirements which must be fulfilled by all materials are corrosion resistance,
biocompatibility, bio-adhesion (bone in growth), bio-functionality (adequate mechanical
properties, especially fatigue strength and a Young's modulus as close to that of the bone as
possible), process ability and availability. These requirements are more or less satisfactorily
fulfilled by the various customary groups of orthopaedic materials [33]. Compared with other
biomaterials like ceramics and polymers, the metallic biomaterials possess the outstanding
property of being able to endure tensile stresses, which, in the case of alloys, may be extremely
high and also of dynamic nature. Typical examples for such highly loaded implants are hip and
knee implants, plates, screws, nails, and dental implants. In comparison different materials show
a different behavior according to the demands. A corrosion resistant material may not necessarily
be biocompatible and, contrarily, a more biocompatible material may be less corrosion resistant.
Presently the typical metallic biomaterials used for implant devices are [34]:
− 316L stainless steels
− Cobalt-chromium alloys
− Ti-6A1-4V alloys
− Ni-Ti alloys
The materials were originally developed for industrial uses. They were subsequently used
in many implant devices, as a biomaterial, due to their relatively high corrosion resistance. The
mechanical properties of these metals, alloys and bone as recommended by ASTM are given in
12
Table 2.1. These materials are accepted by the body environment because of their passive and
inert oxide layer formed on the surface. The main elemental constituents, as well as the minor
alloying constituents of these materials are usually tolerated by the body in trace amounts, since
most of these alloying elements have specific biological role [27].
TABLE 2.1
MECHANICAL PROPERTIES OF METAL ALLOYS AND BONE
Tensile
Yield
Young’s
Fatigue
Strength
Strength
Modulus
Limit
(MN/m)2
(MN/m)2
(GN/m)2
(GN/m)2
316 L SS (Annealed)
650
280
211
0.28
Wrought Co-Cr Alloy
1540
1050
541
0.49
Cast Co-Cr Alloy
690
490
241
0.30
Titanium
710
470
121
0.30
Ti-6Al-4V
1000
970
121
_
Human Bone
137.3
_
30
_
Material
2.1.6 Stainless steel
The austenitic steels, especially Types 316 and 316L, are most widely used for implant
fabrication. It is comprised of Fe 60-65 wt%, Cr 17-19 wt % and Ni 12-14 wt%. Stainless steel
that has a low content of impurities and a passivated finish is entirely suitable for implantation in
the human body. Forged stainless steel has greater yield strength than cast stainless steels, but
13
has lower fatigue strength than other implant alloys. However stainless steel is more ductile and
more easily machined and recent advancements have significantly enhanced its properties.
Because a femoral component fracture with early designs, stainless steel is no longer used
routinely, from the standpoint of erosion, biocompatibility, and fatigue life, stainless steel is
inferior to other super alloys [34, 35].
The only difference in composition between 316L and 316-stainless steel (SS) is the
content of carbon. A wide range of properties exists depending on the heat treatment (annealing
to obtain softer materials) or cold working (for greater strength and hardness). Even 316L
stainless steels may corrode inside the body under certain circumstances in a highly stressed and
oxygen depleted region, such as contact under screws of fracture plates. Thus, stainless steels are
suitable to use only in temporary implant devices, such as fractures plates, screws and hip nails.
Stainless steel, mostly of type 316L,is a very common metal for in-vivo applications, but with
today’s metallurgical technology it is recognized as inferior to both titanium and chromiumcobalt alloys [36].
2.1.7 Cobalt Chrome Alloys
These are basically two types of cobalt chromium alloys; one is the cobalt CoCrMo alloy,
which is usually used to cast a product and the other is the CoNiCrMo alloy, which is usually
wrought by hot forging. Both have excellent biocompatibility in the human environment. The
cast-able CoCrMo alloy (vitallium is a CoCrMo alloy) has been used for many decades in
dentistry and recently in making artificial joints due to corrosion resistant capability. The
wrought CoNiCrMo alloy is a relative newcomer which is used for making the stems of
prosthesis for heavily loaded joints such as the knee and hip. This alloy has a higher degree of
corrosion resistance in salt water when under stress. It has a higher fatigue and ultimate tensile
14
strength than CoCrMo. Over all Cobalt-Chrome alloys are good for components with long
service life requirements. Both groups of alloys contain more than 20 wt% chromium, thus
providing a good resistance due to a passive oxide layer on the surface [37].
Cobalt chrome alloys are wear resistant and heat resistant. They maintain their strength
even strength even at high temperatures. Most of the mechanical properties of these alloys arise
from the crystallographic nature of cobalt while the presence chromium increases corrosion
resistance. Cobalt-based alloys are highly resistant to corrosion and especially to attack by
chloride within crevice. As in all highly alloyed metals in the body environment, galvanic
corrosion can occur, but to a lesser extent than in the iron-based alloys. Cobalt-based alloys are
quite resistant to fatigue and to cracking caused by corrosion, and they are not brittle, since they
have a minimum of 8% elongation. However, as is true of other alloys, cobalt based alloys may
fail because of fatigue fracture but less often than stainless steel stems [36]. Cobalt-based alloys
are not brittle and are resistant to corrosion, however like other alloys they may fail under fatigue
fracture .Even if the ultimate tensile strength of the cobalt-based alloys changes, which could be
caused by cyclic loading or corrosion, the elastic modulus of the material remains constant .The
modulus for these materials are higher than other materials such as stainless steel, however the
affects of this modulus on fracture fixation are not known at this time.
2.1.8 Titanium and Titanium Based Alloys
Attempts to use titanium for implants fabrication dates to the late 1930’s .It was found
titanium was tolerated in cat femurs, as was stainless steel and CoCrMo alloy. Titanium’s
lightness and good mechanical-chemical properties are salient features for implant application.
One titanium alloy (Ti6A14V) is widely used to manufacture implants. The main alloying
15
elements of the alloy are aluminum-Al (5.5-6.5 wt %) and vanadium-V (3.5-4.5 wt %). It is 60%
lighter than Stainless Steel but more expensive [22].
Whilst the strength of the titanium alloys varies from lower than to equal to that of 316
stainless steel, when compared by specific strength (strength per density), the titanium alloys
excel any other implant material. It is a high mechanical resistance alloy which has highest
strength-to-density ratio of all metals up to 500 C. Titanium nevertheless, has poor shear
strength, making it less desirable for bone screws, plates and similar applications. Titanium also
tends to gaul or seize when in sliding contact with itself or other metal. It is brittle i.e. less ductile
than stainless steel, but more ductile titanium alloys are being produced. However it can be as
strong as stainless steel [34, 35].
Titanium-based alloys that have a high co-efficient of friction can cause problems. It has
poorer wear characteristics than other alloys. Wear particles are formed in a piece of bone if a
piece of bone rubs against the implants, or if two parts of an implant rub against one another.
Therefore, implants of titanium upon titanium generally are not used as joint surfaces or load
bearing surface. A downside to Titanium is it has poor shear strength, which limits its application
in bone screws, plates and other fracture fixation devices. Since the Young’s modulus is
approximately half that of stainless steel, therefore less risk of stress protection of bone, stress
riser at end of plate or nail [36, 37]. Table 2.2 shows the comparison between the mechanical
characteristics of Stainless Steel, Cobalt Chrome and Titanium alloys. Each alloy has different
characteristics and is unique in its capabilities.
16
TABLE 2.2
MECHANICAL CHARACTERISTICS OF ORTHOPAEDIC ALLOYS
Characteristics
S-Steel
Cobalt-Chrome
Titanium
Stiffness
High
Medium
Low
Strength
Medium
Medium
High
Low
Medium
High
Low
Medium
High
Corrosion
Resistance
Biocompatibility
2.1.9 New Materials for Knee Implants
The durability of knee implants become a serious concern when growing numbers of
physically active “baby boomers” began to develop arthritis while in their 40s and 50s. The pain
and swelling, which could be disabling, make even normal activities such as walking and
climbing stairs an ordeal. Since total knee replacement hit the mainstream in the 1970s, many
implant materials have been tried and abandoned. The most successful has been metal alloys,
combined with a medical grade plastic. Earlier knee replacements, coupled with our tendency to
live longer, have intensified the search for implant materials that won’t wear out. A 15 to 20 year
life span for a knee implant isn’t good enough anymore but has to work for at least 30 years [38].
More recently, ceramics have demonstrated great promise for replying metals in total
joints with the chief benefit of ceramics is their superior wear properties. Not only is the material
much more resistant to abrasion, but it also is kinder to plastic surface with which it combines.
This means the joints last longer. The U.S. Food and Drug Administration recently approved a
17
durable zirconia/zirconium knee implant, popularly known as a “ceramic” knee .It consists of a
metal compound of zirconium covered with a ceramic surface that promises a longer life implant
[39].
Smith & Nephew has now developed the first knee implants made with a ceramic
surfacing technology that creates a zirconia surface over a zirconium substrate. This hybrid
material, called “oxidized zirconium,” pairs the mechanical properties of a metal with the wearfighting capabilities of a ceramic. Neither a coating nor a through and through ceramic, oxidized
zirconium comes from a high temperature oxidation process that changes the surface of wrought
zirconium parts into zirconia. The ceramic zone extends about five microns below the surface.
For the next few microns, the process leaves a gradient of oxygen-enriched metal, which
ultimately gives way to unadulterated zirconium alloy [40]. Smith & Nephew’s ceramics
technology addresses the growing need for longevity by addressing the wear that typically occurs
as metal femoral components slide on the tibia bearing surface made from ultra–high–molecular–
weight–high–density polyethylene (UHMWPE). Over time, metal alloys such as cobalt chrome
develop tiny scratches from abrasive and oxidative wear, roughening their surface just enough to
eat away the polyethylene bearing. According to Gordon Hunter, a materials engineer for Smith
& Nephew a single scratch 2µm deep, with 1µm adjacent peak height, on a metal counter face
can cause a dramatic increase in the wear rate of UHMWPE. Hunter cities three ways in which
oxidized zirconium’s ceramic surface targets wear. For one, the ceramic surface slides with less
resistance. Its coefficient of friction on polyethylene is less than half that of cobalt chrome. For
another, the material resists abrasion. In pin abrasion tests against acrylic bone cement, which
can produce debris in the joint, the oxidized zirconium exhibited 4,900-times the abrasion
resistance of cobalt chrome. Finally, it’s more than as hard as cobalt chrome, giving it greater
18
immunity to scratches. Patients can have replacement surgery well before the age of 65. For
active patients as this material demonstrates such low wear rates. It is also one of the most
biocompatible metals known to man.
2.2 The Knee
The knee is the largest and one of the most complex parts of the body. Almost perfect
knee function is required to perform daily activities. The knee movements include bending,
straightening and rotational motions which combine together to perform routine activities.
Therefore it is more likely to be injured than any other joint in the body.
2.2.1 Anatomy
It is the meeting place of two important bones in the leg, the femur (the thighbone) and
the tibia (the shinbone). The fibula is the smaller shin bone, located next to the tibia. The patella
(otherwise known as knee cap) is made of bone and sits in front of the knee. These four bones
are connected by muscles, ligaments, and tendons which help control motion, provide stability
and brace the joint against abnormal types of motion. The long thigh muscles give the knee
strength. Other parts of the knee, like cartilage, serve to cushion the knee or help it absorb shock
during motion [40]. Different bones of the knee are shown below in Figure 2.2.1.
19
Figure 2.2.1. Bones of a Knee [40]
The knee is a "hinge type" joint which is formed when the lower end of the femur, rotates
on the upper end of the tibia, and the patella, slides in a groove on the end of the femur as the
knee bends and straightens [41]. Two round knobs called femoral condyles are found on the end
of the femur. These condyles rest on the top surface of the tibia. This surface is called the tibial
plateau. The outside half (farthest away from the other knee) is called the lateral tibial plateau,
and the inside half (closest to the other knee) is called the medial tibial plateau. The patella glides
through a special groove formed by the two femoral condyles called the patellofemoral groove.
The knee components can be seen in Figure 2.2.2. Fibula, the smaller shin bone never enters the
knee joint. It does have a small joint that connects it to the side of the tibia. This joint normally
moves very little.
20
Figure 2.2.2. Knee Components [42]
The joint surfaces where the three bones (Femur, Tibia and Patella) touch are covered
with articular cartilage. It covers the ends of the femur, the femoral groove, the top of the tibia
and the underside of the patella. This material is about 1/4 of an inch thick on the patella and 1/8
of an inch thick on the femur and tibia. It is white, shiny with a rubbery and slippery consistency
[43]. This characteristic allows the surfaces to slide against one another without damage to either
surface. The function of articular cartilage is to absorb shock and provide an extremely smooth
surface to facilitate motion. An x-ray of the knee as shown in Figure 2.2.3, normally shows space
(the "joint space") between the femur and the tibia as well as between the femur and the patella.
This is not an empty space but represents the cartilage which does not show up on x-rays as
shown in the Figure 2.2.3.
21
Articular Cartilage
Figure 2.2.3. Articular Cartilage [43]
The ends of knee bones are kept together by strong bands of tissue called ligaments.
There are 4 main ligaments in the knee. On the inner (medial) side of the knee is the medial
collateral ligament (MCL) and on the outer (lateral) side of the knee is the lateral collateral
ligament (LCL). Inside the knee joint, two other important ligaments stretch between the femur
and the tibia: the anterior cruciate ligament (ACL) in front, and the posterior cruciate ligament
(PCL) in back as shown in Figure 2.2.4. Smaller ligaments help hold the patella in the center of
the femoral groove.
Figure 2.2.4. Ligaments in the knee [40]
22
The MCL and LCL prevent the knee from moving too far in the side-to-side direction.
The ACL and PCL control the front-to-back motion of the knee joint. The ACL keeps the tibia
from sliding too far forward in relation to the femur. The PCL keeps the tibia from sliding too far
backward in relation to the femur [40]. This is illustrated in Figure 2.2.5. Working together, the
two cruciate ligaments control the back-and-forth motion of the knee. The ligaments, all put
together, are the most important structures controlling stability of the knee by connecting bones
to bones. Without strong, tight ligaments to connect the femur to the tibia, the knee joint would
be too loose.
Figure 2.2.5. Movement controlled by ligaments [42]
Two special types of ligaments called menisci sit between the femur and the tibia. These
structures are sometimes referred as the cartilage of the knee and act as "cushions" or "shock
absorbers". The menisci differ from the articular cartilage that covers the surface of the joint.
They also help provide stability to the knee. There is a medial meniscus and a lateral meniscus as
shown in Figure 2.2.6.
23
Figure 2.2.6. Menisci of the knee [40]
The two menisci of the knee are important for two reasons: (1) they work like a gasket to
spread the force from the weight of the body over a larger area, and (2) they help the ligaments
with stability of the knee [41]. The menisci actually warp around the round end of the femur to
fill the space between it and the flat shinbone. The menisci are thicker around the outside, and
this thickness helps keep the round femur from rolling on the flat tibia. Without the menisci, any
weight on the femur will be concentrated to one point on the tibia. Weight distribution by the
menisci is also important because it protects the articular cartilage on the ends of the bones from
excessive forces. Any concentration of force into a small area on the articular cartilage can
damage the surface, leading to degeneration over time [40].
Tendons are similar to ligaments, except that tendons attach muscles to bones. The
largest tendon around the knee is the patellar tendon. This tendon connects the patella to the
tibia. This tendon covers the patella and continues up the thigh. There it is called the quadriceps
tendon since it attaches to the quadriceps muscles in the front of the thigh. This is shown in
24
Figure 2.2.7. The hamstring muscles on the back of the leg also have tendons that attach in
different places around the knee joint.
Figure 2.2.7. Knee Tendons [44]
Another part of the knee system is a bursa which is a small fluid filled sac that decreases
the friction between two tissues. Bursae also protect bony structures. There are many different
bursae around the knee but the one that is most commonly injured is the bursa in front of the
patella [43], the prepatellar bursa as shown in Figure 2.2.8.
Figure 2.2.8. Bursae of Knee Joint [45]
25
The extensor mechanism is the motor that drives the knee joint and allows the body to
walk. It sits in front of the knee joint and is made up of the patella, the patellar tendon, the
quadriceps tendon, and the quadriceps muscles. The quadriceps muscles are the main muscles
that straighten the knee. The hamstring muscles are the main muscles that bend the knee [46].
This complete cycle is shown in Figure 2.2.9.
Figure 2.2.9. Walk mechanism [46]
The knee is the largest and most complicated joint which is used for everything from
standing up, sitting, to walking, and running. The knees are the most easily injured part of the
body. It's a weight-bearing joint that straightens, bends, twists and rotates. All this motion
increases the risk of acute or overuse knee injuries. Therefore knee problems are a fairly
common complaint among people of all ages.
2.2.2 Knee Problems
There are many diseases and types of injuries that can affect the knee. In 2003 about 19.4
million visits were made to physicians' offices because of a knee problem [47]. It was the most
common reason for visiting orthopaedic surgeons. Knee problems can be a result of disease or
injury. A number of diseases can affect the knee (arthritis, osteoarthritis, rheumatoid arthritis,
26
and gout). Knee injuries can occur as the result of a direct blow or sudden movements that strain
the knee beyond its normal range of motion. Sometimes knees are injured slowly over time.
Knee problems can also be the result of a lifetime of normal wear and tear.
2.2.3 Knee Diseases
There are many diseases that can affect the knee. The most common disease is arthritis.
Arthritis is a Greek word which literally means inflammation of the joint. However the term is
used loosely to describe any condition in which there is damage to the joint, even in cases where
the inflammation is absent. The most common forms of arthritis are osteoarthritis (OA) and
rheumatoid arthritis (RA) [48].
Osteoarthritis (OA) is by far the most common form of arthritis. More than 10 million
Americans have osteoarthritis of the knee [47]. It begins with the gradual breakdown of joint
cartilage, resulting in pain and stiffness. In this disease, the cartilage gradually wears away and
changes occur in the adjacent bone the joint may lose its normal shape. With further cartilage
breakdown, the ends of the bones grind against one another causing pain. [49]. Figure 2.2.10
shows the illustration for OA.
Figure 2.2.10. Osteoarthritis Knee Joint [44]
27
Osteoarthritis may be caused by joint injury or being overweight. It is associated with
aging and most typically begins in people age 50 years or older. A young person who develops
osteoarthritis typically has had an injury to the knee or may have an inherited form of the disease
[50]. OA can be distinguished into two forms: Idiopathic OA where no one knows for sure what
causes the destruction of the joint cartilage and Secondary OA where the damage to the cartilage
in the knee joint has a known cause such as previous fracture through the joint surfaces or
damage to the meniscus or ligaments of the knee joint [51].
Rheumatoid arthritis, which generally affects people at a younger age than osteoarthritis,
is an autoimmune disease. This means it occurs as a result of the immune system attacking
components of the body. In rheumatoid arthritis, the primary site of the immune system’s attack
is the synovium, the membrane that lines the joint. This attack causes inflammation of the joint.
It can lead to destruction of the cartilage and bone and, in some cases, muscles, tendons, and
ligaments as well as shown in Figure 2.2.11. It also affects other parts of the body including the
blood, the lungs, and the heart [52]. According to National Institutes of Health RA affects over 2
million Americans, or about 1% of the adult population in the United States. This disease is 2 to
3 times more common in women than in men, and generally affects people between the ages of
30 and 60.
28
Figure 2.2.11. Rheumatoid Arthritis Knee Joint [42]
Other knee disease include gout which results from deposits of needle-like crystals of
uric acid in the connective tissue near joints and/or in the joint space which lead to inflammation,
swelling, and pain [49]. Bursitis involves the inflammation of the bursae, which are small, fluidfilled sacs in knee joints. The inflammation may result from arthritis in the joint, or from injury
or infection of the bursae. Symptoms include pain, tenderness, and limited movement of the
involved joint [53]. Infectious arthritis is a form of joint inflammation that is caused by
infectious agents, such as bacteria or viruses. It does not usually last a long time if it is treated
early [54].
2.2.4 Knee Injuries
Mechanical knee problems such as a direct blow or sudden movements that strain the
knee cause knee injuries. Injuries are also caused by wear and tear of knee parts. Knee diseases
like osteoarthritis, and rheumatoid arthritis can also permanently damage knees.
Cartilage Injuries and Disorders
29
Chondromalacia also called chondromalacia patella, refers to softening of the articular
cartilage of the kneecap (Figure 2.2.12). This causes the patella to rub against the lower end of
the femur instead of gliding smoothly across it. This can be caused by injury, overuse, or muscle
weakness, or if parts of the knee are out of alignment. Chondromalacia can develop if a blow to
the knee cap tears off a piece of cartilage or a piece of cartilage containing a bone fragment. This
eventually causes damage to the cartilage. The disorder is common in runners and is also seen in
skiers, cyclists, and soccer players [50].
Figure 2.2.12. Chondromalacia Patella [46]
The meniscus is easily injured if the knee is twisted while bearing weight. A partial or
total tear may occur (Figure 2.2.13). If the tear is tiny, the meniscus stays connected to the front
and back of the knee. If the tear is large, the meniscus may be left hanging by a thread of
cartilage. The seriousness of the injury depends on the location and the size of the tear.
30
Figure 2.2.13. Meniscus Tear [46]
Ligament Injuries
Two commonly injured ligaments in the knee are the anterior cruciate ligament (ACL)
and the posterior cruciate ligament (PCL). The ACL is the most often stretched or torn by sudden
changing motion, jumps or direct contact as it prevents the shinbone from sliding forwards
beneath the thighbone. PCL injuries which do not occur as frequently as ACL injuries usually
result by a direct impact, such as in an automobile accident or football tackle [41]. Figure2.2.14
shows the ACL and PCL ligament injuries.
Figure 2.2.14. ACL and PCL ligament [42]
31
The medial and lateral collateral ligaments are usually injured by a blow to the outer side
of the knee. This can stretch and tear a ligament (Figure 2.2.15). These blows frequently occur in
sports such as football or hockey.
Figure 2.2.15. Medial and Lateral collateral ligament injuries [42]
Tendon Injuries
Tendon injuries range from tendinitis (inflammation of a tendon – Figure 2.2.16) to a
ruptured (torn) tendon. Torn tendons most often occur from overusing a tendon (particularly in
some sports). The tendon stretches like a worn-out rubber band and becomes inflamed. Tendons
can also tear if thigh muscles contract. This is most likely to happen in older people with weak
tendons [41]. One type of tendinitis of the knee is called jumper’s knee. In sports that require
jumping, such as basketball, the tendon can become inflamed or can tear.
32
Figure 2.2.16. Tendinitis Source: Nucleus Medical Art, Inc.
2.2.5 Knee Problem Treatment
Injuries and diseases of the knees are usually treated by an orthopedist. Patients are
recommended a wide range of treatments. However, the effectiveness of different treatments
varies from person to person. The purpose of treatment is to reduce pain, increase function and
generally reduce symptoms. Treatment options can be non-surgical or surgical [41]. Nonsurgical treatments fall into four major groups:
− Health and behavior modifications, such as physical therapy and exercise, weight loss and
education
− Drug therapy, such as pain relievers or COX-2 inhibitors (drugs that interrupt the cycle of
inflammation)
− Intra-articular (within the joint) treatments, such as injections
− Alternative therapies such as herbal remedies, acupuncture or magnet therapy
If non-surgical treatments do not relieve pain and improve function, physicians may
recommend surgery. About 1 on 4 people will require surgical treatment [41]. The purpose of
33
surgical treatments is to reduce pain, increase function and improve overall symptoms. Surgical
treatments include:
− Arthroscopy is a surgical procedure in which small fiber optic telescope is inserted into a
joint through small incisions. The join space is filled with fluid so the surgeon can clearly see
the components. The most common types of arthroscopic surgery include removal or repair
of a torn meniscus, ligament reconstruction, removal of loose debris, and trimming damaged
cartilage [58].
− Osteotomy ("bone cutting") is a procedure in which a surgeon removes a wedge of bone near
a damaged joint. This shifts weight from an area where there is damaged cartilage to an area
where there is more or healthier cartilage [55].
− Arthroplasty is knee replacement procedure. The damaged bone and tissue is replaced by an
artificial joint. If the damage is limited to one side of the knee then a unicompartmental knee
Arthroplasty is performed. If both sides of the knee are affected, a total joint replacement is
performed. The replacement parts are made of metallic alloys and smooth, wear-resistant
plastic (polyethylene) [41].
2.3 Knee Replacement
Patients are recommended for knee replacement (Arthroplasty) when injuries and
diseases lead to joint damage and persistent pain does not improve. Even though total knee
replacement in some form has been practiced for over 50 years, the complexities of the knee
joint only began to be understood 30 years ago [56].
34
2.3.1 Brief History
Hinge Arthroplasty which was created over 100 years ago has been refined over the
years, its succession of increasingly complex designs leading to the present rotating hinge
prosthesis. In 1890 Theophilus Gluck designed and implanted total knee joints made of ivory,
stabilizing these with plaster of Paris and colophony. Early hinge replacements frequently failed
due to poor metallurgy, improper fixation, and frequent infection. The Walldius hinge, made of
acrylic, was first introduced in 1951 and was upgraded to Co-Cr in 1958. This hinge was used
until the early 1970s [57].
A parallel line of development occurred with total knee replacement that was occurring
with total hips. The modern total hip replacement was invented in 1962 by Sir John Charnley, an
orthopaedic surgeon working in a small country hospital in England [62]. A Canadian
orthopaedic surgeon Frank Gunston working with Sir John Charnley applied the principles of hip
replacement to the knee in late 1960’s. He developed a metal on plastic knee replacement
secured to the bone with cement. This was the first metal and plastic knee and the first with
cement fixation (1968). In 1972 John Insall, M.D., designed what has become the prototype for
traditional total knee replacements. This was a prosthesis made of three components which
would resurface all three surfaces of the knee - the femur, tibia and patella (kneecap). They were
each fixed with bone cement and the results were outstanding. This was the first total knee
complete with specific instrumentation to help with accurate bone cutting and implantation.
Since then many designs have been introduced in the unicompartmental knee replacement to
total knee replacement.
Knee surgery for has become fairly routine and is successful around 96% of the time.
One of the first American surgeons to perform this type of knee surgery was Charles O. Bechtol.
35
He started a total hip replacement program in 1969 later designed a knee replacement system
which was widely used and accepted in the U.S. A major improvement was also the development
of accurate instrumentation for installing the new knee surfaces. In 2003, there were 418,000
total knee replacements performed, primarily for arthritis [59]. Current research with knee
replacement is directed at refining the design to improve patient function. The aim is to achieve
greater knee motion and functionality in knee design and replacement procedures so the
damaged knee resembles normal knees.
2.3.2 Knee Implants
Knee implants are designed to reduce pain and improve overall functionality for the
patient. Physicians consider gender, occupation, disability level, pain intensity, interference with
lifestyle and other medical conditions in selecting the appropriate prosthesis for the patient. In
order for the design to work, it should be able to satisfy certain guidelines of rotations and
displacements in order to perform daily activities. Knee motions in various activities have been
measured and it has been concluded that knee prosthesis should allow an internal/external
rotation of -12° to +12° and an anterior-posterior displacement of 13 mm [60]. A maximum
flexion angle of 125° is advised [61].
The first implant designs used the hinge concept and literally included a connecting hinge
between the components. Newer implant designs, recognizing the complexity of the joint,
attempt to replicate the more complicated motions and to take advantage of the posterior cruciate
ligament (PCL) and collateral ligaments for support. Up to three bone structures can be replaced
during a knee replacement surgery: the lower ends (condyles) of the femur, the top surface of the
tibia and the back surface of the kneecap as shown in Figure 2.3.1 [41]. Components are
36
designed so that metal always articulates against plastic, which provides smooth movement and
results in minimal wear.
Figure 2.3.1. Components of a knee implant [41]
The metal femoral component curves around the end of the thighbone and has an interior
groove so the kneecap can move up and down smoothly against the bone as the knee bends and
straightens. Usually, one large piece is used to resurface the end of the bone. If only one side of
the thighbone is damaged, a smaller piece may be used (unicompartmental knee replacement) to
resurface just that part of the bone. Some designs (posterior stabilized designs) have an internal
post with a circular-shaped device (cam) that works with a corresponding tibial component to
help prevent the thighbone from sliding forward too far on the shinbone when you bend the knee.
The tibial component is a flat metal platform with a polyethylene cushion. The cushion
may be part of the platform (fixed) or separate (mobile) with either a flat surface (PCL-retaining)
or a raised, sloping surface (PCL-substituting).Patellar component: The patellar component is a
37
dome-shaped piece of polyethylene that duplicates the shape of the kneecap anchored to a flat
metal plate.
There are several knee implant designs available in the clinical market today. These
include the fixed bearing knees, mobile bearing, unicompartmental knees, cemented or
cementless among others. Some of the designs are explained below.
Fixed Bearing
The main advantages of a fixed-bearing knee is its simplicity and, as the clinical
experience has shown, its reliability with good to excellent function in the large majority of cases
and a survivorship of around 95 percent at 10 years [62]. The fixed bearing (PCL sacrificing and
PCL retaining) prosthesis can be seen in Figures 2.3.2. Most people get a fixed-bearing
prosthesis that reduces knee pain dramatically and may last for many years.
Figure 2.3.2. Fixed Bearing Knee Model [63]
Mobile Bearing
Compared to the fixed bearing knee implants, mobile bearing knee implants are relatively
new in the surgical market. A major advantage of the mobile-bearing implants is conformity,
which is achieved without the constraint that can lead to excessive forces on the fixation surface
38
[64]. Figure 2.3.3 shows the mobile-bearing prosthesis. This knee design is less forgiving of
imbalance in soft tissues. Bearings without stops can dislocate or spin-out and bearings with
stops can wear against the mechanical stops that prevent dislocation [65].
Figure 2.3.3. Mobile Bearing Knee Design [42]
Unicompartmental
A partial knee replacement - also called a unicompartmental knee replacement -- involves
putting an implant on just one side of the knee, rather than over the entire surface of the knee
joint (Figure 2.3.4). A unicompartmental knee replacement is done if part of the knee joint is
damaged by arthritis and the other compartments have healthy, normal cartilage [41].
39
Figure 2.3.4. Unicompartmental Knee Design [63]
PCL-Retaining
Almost two-thirds of the knees used world wide are of the posterior cruciate ligament
(PCL)-Retaining type, although there is a gradual trend towards the posterior-cruciatesubstituting designs [62]. In PCL-Retaining designs, rearward movement of the tibia is resisted
by an intact PCL, which creates stability (Figure 2.3.5).
Figure 2.3.5. PCL-Retaining Knee Design [63]
40
PCL-Substituting
PCL-Substituting knees (also called posterior stabilized knees) have a raised sloping
surface or a polyethylene post that compensates for the missing PCL to give the knee more
stability [63] (Figure 2.3.6).
Figure 2.3.6. PCL-Substituting Knee Design [63]
Cemented Fixation
Knee replacements may be “cemented” or “cementless” depending on the type of fixation
used to hold the implant in place. The majority of knee replacements are generally cemented into
place. Cemented knee replacements have been used successfully in all patient groups for whom
total knee replacement is appropriate, including young and active patients with advanced
degenerative joint disease. 15 years of clinical reports support this conclusion [41].
Cementless Fixation
Implant designs were introduced in the 1980's that were intended to attach directly to the
bone without the use of cement. Cementless designs rely on bone growth into the surface of the
implant for fixation. Most implant surfaces are textured or coated so that the new bone actually
grows into the surface of the implant. Screws or pegs may also be used to stabilize the implant
41
until bone ingrowth occurs. Because they depend on new bone growth for stability, cementless
implants may require a longer healing time than cemented replacements [41].
2.4 Customized Knee Implant Design
Harrysson [2] proposed a proof-of-concept to develop customized implants for young
patients that minimized bone removal and maximized the fit and usability of the implant [70].
The research emphasized on the fact that even though more knee implant surgeries are taking
place, traditional knee implants are still not able to meet the requirements. As discussed early in
the literature that all knee implants are based on the same basic design. His literature research
concluded that one of the main problems with traditional knee design is the longevity of the
implant. Due to this patients have to go through several revision surgeries throughout their life
time. Patient’s bones have to be customized to fit the implant and each time a revision is
performed more bone is removed. Since the bone is cut to match the implant, inmost cases the fit
is not perfect which leads to loosening of the implant. Implants also loose due to higher activity
level. This limits the implant to patients who are younger than 60 years. Younger patients have a
lower success rate with traditional knee implant design. One of the aims of his research was to
develop a customized knee implant that can also be used by younger patients.
The proposed research to develop a customized implant was targeted towards the younger
patients in need of Total Knee Arthroplasty. Older patients can not benefit from this proposed
methodology due to economic and practical reasons. The proposed methodology for
customization and optimization of the knee implant was broken down into the following steps:
1. Acquisition of patient’s knee CT-scan data
2. Conversion of CT-data to a usable CAD-model
3. Design of the implants using CAD-model as base
42
4. Use Rapid Tooling to produce a cast master of the implant components
5. Machine the components to final finish and application of porous coating
6. 3D laser scanning of components for accurate 3D surface representation
7. Programming of the orthopaedic robot based on the scan data
The femur component design was started by acquiring the CT-scan data of Ola’s due to
lack of public available patient data. The CT-scan was performed at the Center of Diagnostic
imaging in Florida on a GE scanner. The CT-image (Figure 2.4.1) was stored on an optical disc
and transferred to the lab where it was converted into a 3D computer model using Mimics (3D
imaging software from Materialise, Belgium).
Figure 2.4.1. CT-images of author’s knee. [2]
Due to the file format limitations in Mimics, 3D model of distal femur (Figure 2.4.2) was
not successfully imported into a CAD package.
43
Figure 2.4.2. 3D model of distal femur [2].
After several attempts using Pro-E, it was concluded that a solid CAD-model of the
femur was not necessary for the custom design of the implant component. It was also realized
that a solid CAD model would not provide sufficient information and a better way of
representing the patient’s femur geometry was required.
Finally Magics, software from Materiliase-Belgium, was selected to transfer the cross
sections of 3D knee model. The cross sections were exported as iges-curves for an exact and
precise representation of the knee surface. The iges-curves were imported into Pro-E and a wire
frame model representing the distal portion of the femur was created as shown in Figure 2.4.3.
44
Figure 2.4.3. Iges-curves imported into Pro-E [2]
The imported curves in Pro-E could not be used directly to create features so new curves
were created on top of the imported curves. Splines were used to create the best possible result
which was time consuming but it helped in correcting any imperfections in the surface of the
distal femur caused by either data loss within the conversion process or actual fractures or
damages on the bone. A solid model of the outer configuration was created by blending the
curves together as shown in Figure 2.4.4.
Figure 2.4.4. Spline curves created in Pro-E [2]
45
The interface surface between the implant and the bone structure was designed next to
solve the problem with premature loosening of the component. The goal was to preserve as much
of the natural bone as possible and make the implant as thin as possible without risking
mechanical failure of the component. The internal surface was created using the same datum
planes as the outer surface. The shapes of the curves were based on the designer’s intention after
several iterations. Additional cuts and changes are made to achieve the final desired design,
shape and natural curvature as shown in Figure 2.4.5). The tibia component was also designed
similarly to achieve the customized requirements.
Figure 2.4.5. Customized Femoral Design [2].
46
CHAPTER 3
COMPARISON OF KNEE DESIGNS
The objective of this research was to compare the femoral component of customized knee
design and traditional knee design with femur bone interface. The research was under taken to
fill the gap in Harrysson’s [2] work. His research did not offer a reliable comparison of the
traditional and customized knee designs. The comparison was performed using simple load
conditions and without femur bone interface. In addition Harrysson [2] did not consider the
options for possible orthopaedic materials. The generated knee model was also not suitable for
Finite Element Analysis.
3.1 Research Methodology
This research concentrated on the comparison of femoral component for traditional and
customized implants with bone interface. The procedure started with creating the CAD models
for both traditional and customized designs. A general factorial design was set to compare the
knee component models. The factors used in the factorial design included three load levels
applied on the femur bone at two gait knee angles using four orthopaedic materials with the
Maximum Stress Level (
max)
as the measured response. Finite Element Analyses (FEA) was
performed on the knee models to measure the response (Maximum Stress Level -
max).
The
response along with the factors was entered in computer software to perform ANOVA. The
results are used to draw conclusions. The research procedure is outlined in the flow chart in
Figure 3.1.
47
Design a General
Factorial to
compare models
Prepare CAD
models for both
designs
Perform Finite
Element
Analysis
Analyze
Results
(ANOVA)
Draw
Conclusions
Figure 3.1 Research Procedure
3.2 Customized Femoral Component Model
The customized femoral component of the knee implant was provided by Harrysson [2].
The femoral component was customized using the CT-image from his knee. The different stages
in acquiring the CT-image to creating the knee model were explained in Chapter 2. The femoral
component was provided in Initial Graphics Exchange (iges) format. The format primarily
consists of points, lines and surfaces.
Harrysson [2] used Pro-E to convert the iges knee files into a usable format for FEA.
Catia V5 R17 [67] was selected as the tool for CAD model creation from iges file due to limited
experience with Pro-E.
The knee iges file was imported into Catia using built-in functions. The imported file is shown in
Figure 3.2.
48
Figure 3.2. Different View Angles of Imported Knee File in Catia V5 R17 [67]
After initial review, it was found that the imported knee file was made up of points,
curves and surfaces stitched together to form a hollow object. The next step was to convert this
hollow object into a solid model. This process was unsuccessful because the surfaces were not
stitched together but placed very close to each other to form the knee. The healing function was
used to stitch the surfaces together. The object was converted to a solid model for FEA. The file
was converted from iges format to a solid model file (.catpart).
Finite Element Analysis (FEA) software was used to mesh the model and measure
maximum stress level. Catia V5 R17 was selected as the pre and post processor for this task. The
step was unsuccessful due to the unsuitable solid model. After several attempts it was decided to
build a suitable solid model using the outer contour points of the femoral component from the
iges file. Since the objective is to measure the maximum stress level, the CAD model was
49
created using the thinnest area in the knee implant where the stress level is assumed to be at its
peak.
The knee implant was sliced vertically to create cross sections at 7 locations due to the
thickness configuration of the customized implant. The implant with marked cross sections is
shown in Figure 3.3. The thinnest cross-section in the knee implant, which touched the tibia
component, was selected.
Figure 3.3. Sliced cross-sections of femoral component
The cross section provided several points on the outside contour of the knee component.
The points were connected carefully by using splines. Several iterations were made to make sure
the final outline resembled that of the cross section of the femoral component. Once the contour
50
was finalized, it was extruded 0.1 inch to build the solid cross section model of the knee femoral
component as shown in Figure 3.4.
Figure 3.4. Solid Cross Section of Customized Knee Design Model
As discussed in the Chapter 2 the knee implant should allow a maximum flexion angle of
approximately 125º. Therefore it was necessary to make sure that the selected cross section was
thinnest at all the locations which take the maximum loads during walking, running, jumping and
climbing stairs. The extra step was taken because the thickness of the customized implant is not
consistent at the condyle locations (area where the femoral component meets the tibia
component).
It was confirmed that the selected cross section was only thinnest at the 0º angle or when
the patient is standing but not at the 45º angle. Therefore another cross section was selected
51
which represented the thinnest location at 45º angle. The solid femoral knee model was built
from the new cross section. Finally one cross section represented the 0º angle (standing) and the
second represented the 45º angle (walking).
The femur bone was modeled and attached to the implant cross section solid. The contact
between the bone and implant was assumed to represent a perfect fit. The femur bone surface
attached to the implant was built using the inner contour of the implant model. The profile of the
bone part (not in contact with the implant) was extended to resemble the stem of the femur bone.
The final assembly is shown in Figure 3.5.
Figure 3.5. Femur bone attached to the customized knee implant.
52
3.3 Traditional Femoral Component Model
The femoral component for the traditional knee implant model was built in the same
manner i.e. the solid knee model was extruded using the contour built from points and splines.
The outer contour designed was modeled following the prosthetic industry’s standard where the
surface of the femoral component is formed using two circles and taken on an eclipse shape [61]
Figure 3.6. The design process was an approximation and took several iterations to represent the
shape of the industry standard.
Figure 3.6. Industry Design Standard for Femoral Component [61].
The inner contour was modeled after parametric description of the geometry of conventional
knee implants as described by Walker [62] as shown in Figure 3.7.
53
Figure 3.7. Parametric Description of Femoral Component Geometry [62].
The process took several iterations and was good approximation of the standard. Since
the thickness is constant in traditional knee designs, only one cross section was built as shown in
Figure 3.8.
Figure 3.8. Solid Cross Section of Traditional Knee Design Model.
54
The femur bone was modeled and attached to the implant cross section solid. In this case
the contact between the bone and implant was also assumed to represent a perfect fit. The femur
bone surface attached to the implant was built using the inner contour of the implant model. The
profile of the bone part (not in contact with the implant) was extended to resemble the stem of
the femur bone. The final assembly is shown in Figure 3.9.
Figure 3.9. Femur bone attached to the traditional knee implant.
3.4 The General Factorial Design
A General Factorial design was selected to study the joint effect of factors on the
response variable, Maximum Stress Level (
max).
Table 3.1 describes the control factors and their
selected levels.
55
TABLE 3.1
DESIGN FACTORS AND THEIR LEVELS
Factor
Symbol
Design
A
Material
B
Stainless Steel
Angle
C
0º
Weight
D
3*Body Weight
Level
Customized
Traditional
Cobalt Chrome
Ti6Al4V Oxidized Zirconium
45º
5*Body Weight
10*Body Weight
The total number of simulation runs was derived by multiplying the factor levels (2 x 3 x
2 x 4 = 48). The response variable, Maximum Stress Level was measured for 48 simulations. The
design matrix was created in Design Expert [68]. Boundary and load conditions are considered
as held constant factors.
3.5 Finite Element Analyses
Once the customized and traditional femoral component models were built, Catia V5 R17
[67] was used to mesh the solid models and run the FEA simulations. The first step was to apply
the boundary conditions, loads and material properties on the three solid models (2 models
representing the customized design model, 1 representing the traditional design model). Each
design was analyzed to simulate at 2 different gait angles (0º and 45º) representing various
activities. Since the customized design is represented by 2 solid models, the first model was
analyzed at 0º gait angle representing the thinnest structure at this posture. The second model
was analyzed at 45º representing the thinnest structure at this posture. The traditional knee design
model was analyzed to simulate at all the two gait angles.
56
The boundary conditions applied on the implant model and femur bone are illustrated in
Figure 3.10 and Figure 3.11. The contact surfaces between the femur bone and implant are
selected to simulate the link between the two models.
Femur Bone
- Contact Surface
Implant
Figure 3.10. Loads and Boundary Conditions at 0° Angle
Femur Bone
- Contact Surface
Implant
Figure 3.11. Loads and Boundary Conditions at 45° Angle
57
The boundary conditions for the nodes at the contact surface between the condyles and
the tibia bearing plate were restricted to six degrees of freedom. The bone was restricted to single
degree of translation freedom. The loads are applied as a distributed force on the surface of the
femur bone. Three different types of loads were applied at two different gait angles. Harrysson
[2] reported that the knee joint experienced loads equal to 3 times the bodyweight during normal
walking, 5 times the bodyweight during stair stepping and 10 times the body weight during
running or jumping. Table 3.1 shows the loads used in the analysis:
TABLE 3.2
LOAD LEVELS
X-Weight
Patient's Weight (lbs)
Total Weight (lbs)
3
188.5
565.5
5
188.5
942.5
10
188.5
1885
The implants were analyzed using four different materials: Stainless Steel, Cobalt
Chrome, Titanium (Ti6Al4V) and Oxidized Zirconium. Table 3.2 shows the properties of the
materials used in FEA. The femur bone and implant model were meshed using linear tetrahedral
solid elements. Contact surface properties were used to connect and simulate the link between
the implant model and femur bone.
58
TABLE 3.3
MATERIAL PROPERTIES
Material
Young’s Modulus x 106 psi
Poisson Ratio
316 L Stainless Steel
30.0
0.29
Co-Cr Alloy
34.8
0.33
Titanium (Ti-6Al-4V)
16
0.31
Oxidized Zirconium
14.5
0.27
Femur Bone
0.056
0.30
Catia V5 R17 ELFINI [67] was used as the solver for FEA. Three different loads with boundary
conditions were applied on the femur bone at two different gait angles using four material
properties on the customized and traditional design models to simulate different types of activity.
In total 48 simulations were run to measure the response variable (Maximum Stress level using Von Mises stress analysis.
59
max)
Results
The results (Response Variable: Maximum Stress Level, psi) from the analyses on the
customized and traditional designs are shown in Table 3.4. A sample of the FEA results is shown
in Figure 3.12.
TABLE 3.4
DESIGN MATRIX WITH FEA RESULTS (MEASURED RESPONSE)
Factor
Run
1
2
3
4
5
6
7
8
9
10
11
12
13
14
15
16
17
18
19
20
21
22
23
24
25
26
27
28
29
A:Design
B:Material
Customized
Traditional
Traditional
Traditional
Traditional
Traditional
Customized
Customized
Traditional
Customized
Traditional
Customized
Traditional
Customized
Traditional
Customized
Customized
Traditional
Traditional
Customized
Customized
Customized
Customized
Traditional
Traditional
Customized
Customized
Traditional
Customized
Ti6Al4V
Co. Chrome
Ox. Zirconium
Ti6Al4V
Ox. Zirconium
Co. Chrome
Co. Chrome
Ox. Zirconium
Ox. Zirconium
Ox. Zirconium
Ti6Al4V
Ti6Al4V
Ti6Al4V
Ox. Zirconium
Ti6Al4V
Co. Chrome
Co. Chrome
Co. Chrome
Ox. Zirconium
Ox. Zirconium
St. Steel
Ox. Zirconium
Co. Chrome
Ox. Zirconium
St. Steel
Ti6Al4V
St. Steel
Co. Chrome
Co. Chrome
60
C:Angle D:Weight
Degrees
lbs
45
1885
0
942.5
0
565.5
45
1885
0
942.5
45
565.5
45
565.5
0
565.5
0
1885
0
942.5
0
942.5
45
942.5
0
565.5
45
565.5
45
942.5
45
942.5
0
942.5
0
565.5
45
1885
45
1885
0
565.5
0
1885
0
565.5
45
565.5
0
942.5
0
1885
45
942.5
45
942.5
0
1885
Response
Max. Stress
Psi
17632831
148927
89070
8373181
148450
2504786
5265501
93980
296901
156633
148312
8816416
88987
5397967
4186590
8775835
188158
89357
8409400
17993225
108548
313266
112895
2520589
149350
322176
8967580
4178337
376317
TABLE 3.4 (continued)
DESIGN MATRIX WITH FEA RESULTS (MEASURED RESPONSE)
Factor
Run
30
31
32
33
34
35
36
37
38
39
40
41
42
43
44
45
46
47
48
A:Design
B:Material
Traditional
Traditional
Traditional
Customized
Traditional
Customized
Traditional
Customized
Customized
Traditional
Traditional
Customized
Traditional
Customized
Traditional
Traditional
Customized
Customized
Customized
St. Steel
St. Steel
Ti6Al4V
St. Steel
Co. Chrome
Co. Chrome
Ox. Zirconium
Ti6Al4V
St. Steel
St. Steel
St. Steel
Ti6Al4V
Ti6Al4V
St. Steel
St. Steel
Co. Chrome
Ox. Zirconium
Ti6Al4V
St. Steel
C:Angle
Degrees
0
45
0
45
0
45
45
0
0
45
45
0
45
45
0
45
45
45
0
D:Weight
lbs
1885
1885
1885
1885
1885
1885
942.5
565.5
942.5
942.5
565.5
942.5
565.5
565.5
565.5
1885
942.5
565.5
1885
Response
Max. Stress
Psi
298700
8390636
296625
17935159
297856
17551670
4204700
96652
180913
4195318
2514965
161088
2509733
5380548
89610
8356675
8996613
5289849
361828
Figure 3.12. Sample FEA results for Traditional and Customized Implants
61
The results from the FEA show that the Maximum Stress Level (
max)
in customized
implant design is higher than the traditional implant design for all the combinations considered.
The results are compared graphically in Appendix A. Each graph shows the comparison of
Maximum Stress Level (
max)
for both designs for each gait angle, load case and materials.
Overall the stress level is higher in the customized design when compared to the traditional
design for all the four materials. The results of the analyses from Catia V5 R17 [67] are shown in
Appendices B through I, divided according to the four orthopaedic materials.
Since the objective was to compare the effect of different materials in knee implants, it is
observed that the stress level changes with the change in material. However the difference in
max
between the four materials is small. The results (Maximum Stress Level -
max)
from Table
3.4 was used for the analysis of variance in the next section to determine if orthopaedic materials
have significant effect on the knee design with respect to the Maximum Stress Level (
max).
3.4 Statistical Analysis
The Analysis of Variance (ANOVA) was performed on the FEA results (Response
Variable: Maximum Stress Level -
max)
from Table 3.4 using Design Expert [68]. There are
four factors: Design (Factor A), Material (Factor B), Angle (Factor C), and Weight (Factor D).
Maximum Stress Level was measured as the response variable. The resulting ANOVA is shown
in Table 3.5.
62
TABLE 3.5
ANOVA TABLE
Source
Sum of
Squares
DF
Mean
F
Square
Value
Prob > F
Model
174.2596
4
43.56489
5313.503
< 0.0001
A
2.389834
1
2.389834
291.4822
< 0.0001
C
159.3387
1
159.3387
19434.15
< 0.0001
D
11.40883
1
11.40883
1391.507
< 0.0001
AC
1.122213
1
1.122213
136.8736
< 0.0001
Residual
0.352553
43
0.008199
An examination of the residuals indicated no violations of the underlying assumptions of
the ANOVA procedures. The residual plots are shown in Appendix J. The ANOVA indicates a
statistically significant two-factor interaction involving the Design (A) and Angle (C). The
Weight (D), as expected, has a significant effect on the Maximum Stress Level (
max).
The
ANOVA also indicates that orthopaedic materials (B) do not have a significant effect on the
response variable (Maximum Stress Level -
max).
Figure 3.13 shows the design-angle interaction. At higher level of Angle (C), a change in
the Design appears to have a larger effect on the average stress level (
average
max
max).
is 5.03x106 psi for the traditional design whereas the average
At the 45°, the
max
is as high as
1.07x107 psi for the customized design. This represents a two fold increase. Overall increasing
the angle causes an increase in the average maximum stress level with the customized design
indicating higher average maximum stress level (
63
max).
17993225
Max. Stress Level -psi
13517166
9041106
4565047
Customized
Traditional
88987
0.00
11.25
22.50
33.75
45.00
C: Angle (Degrees)
Figure 3.13. Design – Angle Interaction
Weight (D) effect plot is shown in Figure 3.14. As expected increasing the weight (D)
tends to increase the average maximum stress level (
higher average maximum stress level (
max)
max).
The change in weight indicates a
for customized design.
64
Max. Stress Level - psi
19192904.9
14416925.4
9640946.0
4864966.5
88987.0
565.50
895.38
1225.25
1555.13
D: Weight (lbs)
Figure 3.14. Weight Effect Plot
65
1885.00
CHAPTER 4
DISCUSSION & CONCLUSIONS
Harrysson’s [2] proposal to design and customize knee implants for young patients has
promising future. However there are several areas of improvement that require further research
before the concept is efficiently implemented. One of the areas for improvement is the process of
gathering and converting CT-data to a usable CAD model for customization and optimization. A
major portion of this research was involved in converting the actual knee data to a solid model
for FEA. Since the process was not mature, several attempts at various technologies were
undertaken to come up with a final usable format of the model for FEA. This process can be
significantly improved given the advancements in image processing and data conversion.
The load and boundary conditions on both the designs are an approximation and do not
represent the actual load conditions on the knee where several other elements including the tibia,
ligaments, muscles, tendons and patella should be included to provide a balanced movement.
Therefore the stress level on both knee implant components is the closest approximation to the
actual loading and boundary conditions.
The customized implant did not perform better than the traditional implant based on the
assumptions made and loading conditions used in this research. Another area of improvement in
the customization process of the implant is to increase the thickness of the cross section
depending on the patient’s weight. The increase will improve the strength of the knee and reduce
the maximum stress level. When the CAD model is created, each implant is designed and
optimized for individual patient’s knee. The design depends on the disease/injury of the knee.
Currently there are no guidelines or standards to customize the knee implant. This is important as
66
each implant has to withstand the load, avoid failure during activities, and sustain it’s design and
characteristics for the complete life which is anticipated to be longer than traditional knee
implants. Guidelines/standards will help the manufacturers to design and optimize the individual
implant in an efficient manner.
Metal alloys play an important role in defining knee implants. Most metal alloys have
been tested and used for several decades. Each material has unique physical and biological
properties which are important for implants. Surgeons use specific knee designs with specific
materials for each patient. In most cases the knee implant is assigned to the patient based on
his/her medical history, allergies, physical activities and the surgeon’s preference. The
preference is usually based on the success of the implant that relates to the implant’s
biocompatibility, wear and fatigue properties. The results from this research indicated that all the
four orthopaedic materials had roughly the same stress level (
max)
for the same design under
identical load conditions.
The scope of the research was limited to the Stress level (
max)
for each material which
was selected as the deciding factor for comparison between the customized and traditional knee
designs. However the knee designs were not compared using the biocompatibility, wear or
fatigue properties of orthopaedic materials. The literature review has highlighted these properties
as major players in the success or failure of knee implants. Along with the strength, implants
should also be fatigue and wear resistant and should be easily accepted by the human body.
Therefore it could not be confirmed which material is significantly better than the other based on
the factors relating to fatigue, wear and biocompatibility for either customized or traditional
implant.
67
One of the materials (Oxidized Zirconium) considered in this research is relatively new in
the market. Knee implants made from this material have been introduced by Smith & Nephew
[42]. This material has shown promising results due to its unique mechanical properties of a
metal and the wear-fighting capabilities of a ceramic. During FEA it performed similar to the rest
of the materials.
The proposal for customized implants is expected to work if all the processes are
seamlessly integrated and cost efficient. This was confirmed during several conversations with
major manufacturers of knee implants including Zimmer Corporation, Smith & Nephew, and
Depuy Orthopaedics.
There are many advantages to this proposed methodology. It holds promising results
which can lead to fewer revisions and extend the longevity of the implant due to the improved
fit. It can also be used with younger patients with higher activity levels. It allows the surgeons to
remove less bone in order to make the bone fit the implant which is the standard practice today.
Further more the customized implant has not been analyzed using actual knee conditions
and complete knee implant components. It is recommended that further research be done in
streamlining the knee geometry capture process and conversion of this data to a more usable
CAD model process. Efforts should also be made to develop standards for designing the implant
for each patient. FEA analysis should be performed using actual 3D bone model attached to the
complete 3D implant along with other implant components. The concept holds promising future
if all the areas are working efficiently to produce a cost efficient solution.
68
REFERENCES
69
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15,
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APPENDICES
75
APPENDIX A
Maximum Stress Level (
max)
Comparison Charts
0-ANGLE
4.00E+05
3.50E+05
Maximum Stress Level - psi
3.00E+05
St. Steel - Trad
St. Steel - Cus t
2.50E+05
Ti6Al4V - Trad
Ti6Al4V - Cus t
2.00E+05
Co. Chrom e - Trad
Co. Chrom e - Cus t
1.50E+05
Ox. Zirconium - Trad
Ox. Zirconium - Cus t
1.00E+05
5.00E+04
0.00E+00
3
5
10
X - Body W eight
Comparison of Knee Designs at 0º Gait Angle
45-ANGLE
2.00E+07
1.80E+07
Maxim um Stress Level - psi
1.60E+07
1.40E+07
St. Steel - Trad
St. Steel - Cus t
1.20E+07
Ti6Al4V - Trad
Ti6Al4V - Cus t
1.00E+07
Co. Chrom e - Trad
Co. Chrom e - Cus t
8.00E+06
Ox. Zirconium - Trad
6.00E+06
Ox. Zirconium - Cus t
4.00E+06
2.00E+06
0.00E+00
3
5
10
X - Body W eight
Comparison of Knee Designs at 45º Gait Angle
76
APPENDIX B
Finite Element Analysis of Traditional Knee Design (Stainless Steel)
3 times Body weight at 0º Angle
3 times body weight at 45º Angle
77
5 times weight at 0º Angle
5 times body weight at 45º Angle
78
10 times body weight at 0º Angle
10 times body weight at 45º Angle
79
APPENDIX C
Finite Element Analysis of Traditional Knee Design (Cobalt Chrome)
3 times Body weight at 0º Angle
3 times body weight at 45º Angle
80
5 times body weight at 0º Angle
5 times body weight at 45º Angle
81
10 times body weight at 0º Angle
10 times body weight at 45º Angle
82
APPENDIX D
Finite Element Analysis of Traditional Knee Design (Ti6Al4V)
3 times Body weight at 0º Angle
3 times body weight at 45º Angle
83
5 times weight at 0º Angle
5 times body weight at 45º Angle
84
10 times body weight at 0º Angle
10 times body weight at 45º Angle
85
APPENDIX E
Finite Element Analysis of Traditional Knee Design (Oxidized Zirconium)
3 times Body weight at 0º Angle
3 times body weight at 45º Angle
86
5 times weight at 0º Angle
5 times body weight at 45º Angle
87
10 times body weight at 0º Angle
10 times body weight at 45º Angle
88
APPENDIX F
Finite Element Analysis of Customized Knee Design (Stainless Steel)
3 times Body weight at 0º Angle
3 times body weight at 45º Angle
89
5 times weight at 0º Angle
5 times body weight at 45º Angle
90
10 times body weight at 0º Angle
10 times body weight at 45º Angle
91
APPENDIX G
Finite Element Analysis of Customized Knee Design (Cobalt Chrome)
3 times Body weight at 0º Angle
3 times body weight at 45º Angle
92
5 times weight at 0º Angle
5 times body weight at 45º Angle
93
10 times body weight at 0º Angle
10 times body weight at 45º Angle
94
APPENDIX H
Finite Element Analysis of Customized Knee Design (Ti6Al4V)
3 times Body weight at 0º Angle
3 times body weight at 45º Angle
95
5 times weight at 0º Angle
5 times body weight at 45º Angle
96
10 times body weight at 0º Angle
10 times body weight at 45º Angle
97
APPENDIX I
Finite Element Analysis of Customized Knee Design (Oxidized Zirconium)
3 times Body weight at 0º Angle
3 times body weight at 45º Angle
98
5 times weight at 0º Angle
5 times body weight at 45º Angle
99
10 times body weight at 0º Angle
10 times body weight at 45º Angle
100
APPENDIX J
ANOVA Residual Plots
Normal Plot of Residuals
Normal % Probability
99
95
90
80
70
50
30
20
10
5
1
-1.91
-0.86
0.20
1.25
2.30
Studentized Residuals
Residuals vs. Run
Studentized Residuals
3.00
1.50
0.0
-1.50
-3.00
1
10
19
28
Run Number
101
37
46