751
The Journal of Experimental Biology 203, 751–756 (2000)
Printed in Great Britain © The Company of Biologists Limited 2000
JEB2507
LOAD–ELONGATION CHARACTERISTICS OF IN VIVO HUMAN TENDON AND
APONEUROSIS
1Scottish
CONSTANTINOS N. MAGANARIS1,* AND JOHN P. PAUL2
School of Sport Studies, University of Strathclyde, Glasgow G13 1PP, UK and 2Bioengineering Unit,
University of Strathclyde, Glasgow G4 0NW, UK
*Present address: University of Tokyo, Komaba, Department of Life Sciences (Sport Sciences), Komaba 3-8-1, Meguro, Tokyo 153-8902,
Japan (e-mail: costis.maganaris@strath.ac.uk)
Accepted 10 December 1999; published on WWW 26 January 2000
Summary
In the present study, we measured the in vivo
increased curvilinearly from 1.3 to 4 mm and from 3.7 to
load–elongation characteristics of the human tibialis
12 mm, respectively, as a function of dorsiflexion load.
Scaling of the displacements recorded to the resting lengths
anterior tendon and its central aponeurosis. Measurements
(measured over the skin) yielded strain values that
were taken in five men using dynamometry, muscle
increased curvilinearly with load, from 0.8 to 2.5 % in the
electrical stimulation and ultrasonography. Percutaneous
tendon and from 2.1 to 7 % in the aponeurosis. Tendon
tetanic stimulation of the muscle at successive voltages
corresponding to 20, 40, 60, 80 and 100 % of maximum
strain was smaller by between 61 and 64 % compared with
isometric dorsiflexion moment was applied. During
aponeurosis strain at any given contraction level. These
findings are in line with reports from in vitro isolated
electrical stimulation, we recorded the displacements of
material testing and have important implications for
the tibialis anterior tendon origin and its aponeurosis
muscle modelling.
proximal end using B-mode ultrasonography. Aponeurosis
displacement was calculated by subtracting tendon
displacement from the displacement of the aponeurosis
Key words: tendon, aponeurosis, ultrasonography, electrical
stimulation, load–elongation, human.
proximal end. Tendon and aponeurosis displacements
Introduction
Tendinous tissue elasticity has long been studied using one
of the following four methodologies: (i) the alpha method
(Morgan, 1977), (ii) the spindle null-point method (Rack and
Westbury, 1984), (iii) the free vibration method (e.g. Shorten,
1987) and (iv) tensile testing methodologies (e.g. Bennett et
al., 1986). Each of these methods has advantages and
limitations. The alpha method is adaptable under in vivo
conditions (Cook and McDonagh, 1996) but assumes a linear
stiffness increase with force and takes account of the whole
tendinous component, i.e. extramuscular tendon and
aponeurosis. The latter is also the case for the free vibration in
vivo method and the spindle null-point in situ method. In
tensile testing methodologies, the extra- and intramuscular
tendon portions can be tested separately, but preserved or deepfrozen material that may have altered properties (Smith et al.,
1996) has traditionally been used (e.g. Bennett et al., 1986).
B-mode ultrasonography offers the possibility for realistic
in vivo assessment of human tendinous tissue elasticity. It
results in accurate scanning of intramuscular collagenous
tissue (Kawakami et al., 1993; Narici et al., 1996) and allows
the recording of its movement in intact muscle during
contraction (Fukunaga et al., 1996). Selection of appropriate
reference points in the extra- and intramuscular tendon portions
scanned in combination with artificial selective muscle
activation would allow quantification of tendon and
aponeurosis displacements upon control loading. In the present
study, we combined these techniques to estimate the in vivo
load–elongation characteristics of the human tibialis anterior
tendon and its central aponeurosis. Our hypothesis was that,
for a given load application, tendon elongation would be
smaller than that of the aponeurosis.
Materials and methods
Experimental protocol
Measurements were taken in five healthy male volunteers
(age 22±4 years; height 172±4 cm; body mass 73±7 kg; lower
leg length 40±3 cm, means ± S.D.). The study was approved by
the local ethics committee. Isometric ankle dorsiflexion
moments were measured in the prone position on an isokinetic
dynamometer (Lido Active, Loredan Biomedical, Davis,
California, USA) with the knee of the tested leg (the right leg
in all subjects) flexed at 90 °. The ankle joint axis was visually
aligned with the dynamometer axis, and the foot was fixed at
the neutral anatomical position (the sole of the foot at 90 ° to
the tibia). At that ankle angle, no passive forces were detected
752
C. N. MAGANARIS AND J. P. PAUL
by the dynamometer (see also Siegler et al., 1984). To maintain
the ankle joint and dynamoterer axes alignment during
dorsiflexion contraction, we bandaged the foot and the heel
around the dynamometer footplate using Velcro straps, and we
placed a mechanical stop below the knee. This system
prevented any observable movement of the lower extremity
during maximum voluntary isometric dorsiflexion. The
dorsiflexion forces elicited in our experiment were much
smaller than those generated upon maximum voluntary effort
(see below). We therefore assumed that any foot movement in
our measurements would have been negligible.
Contraction forces were elicited by means of percutaneous
electrical stimulation of the tibialis anterior muscle using a
custom-built, high-voltage stimulator. A train of bipolar wave
pulses with a duration of 100 µs was applied at a frequency of
100 Hz for 1 s through two 4 cm×3 cm aluminium foil pads
placed over the motor point area of the muscle. We first
identified the voltage eliciting maximal joint moment (Mmax).
Maximality was indicated by no further increase in the joint
moment recorded with increasing voltage. We then identified
the voltages eliciting joint moments corresponding to 20, 40,
60 and 80 % of Mmax. Surface electromyographic (EMG)
signals from the nearby soleus and peroneous tertius muscles
showed no evidence of current leakage during stimulation of
the tibialis anterior muscle.
A 7.5 MHz linear-array B-mode ultrasound probe (Esaote
Biomedica, Florence, Italy; width resolution 1 mm, depth
resolution 0.62 mm) was secured with adhesive tape onto the
skin, first over the tibialis anterior myotendinous junction
region, approximately 7 cm above the malleoli, and then over
the proximal region of the tibialis anterior muscle,
approximately 7 cm below the knee. In scans recorded at rest,
we identified the tibialis anterior tendon origin and its central
aponeurosis proximal end. Muscle contractions were then
elicited at successive voltages corresponding to 20, 40, 60, 80
and 100 % of Mmax. The displacements of the tibialis anterior
tendon origin and aponeurosis proximal end in the transition
from rest to all contraction level joint moments were recorded
and digitized. The tibialis anterior tendon origin displacement
was considered as the tibialis anterior tendon displacement,
and the aponeurosis proximal end displacement was
considered as the tendon/aponeurosis complex displacement.
Aponeurosis displacement was estimated by subtracting the
tendon displacement from the tendon/aponeurosis complex
displacement (Fig. 1). From each site, we recorded
displacement data from two series of measurements, 2 min
apart, and we used average values for further analysis.
Measurements were taken after the tibialis anterior tendinous
structures had been preconditioned by five series of stimulated
contractions at 20, 40, 60, 80 and 100 % of Mmax, 30 s apart
(see also Bennett et al., 1986). To calculate strain values from
the displacements obtained, we measured the resting lengths of
the tibialis anterior tendon and aponeurosis to the nearest
millimetre over the skin as described by Ito et al. (1998).
Tendon length was estimated as the length from the tibialis
anterior tendon origin to its insertion point on the base of the
first metatarsal bone following the curved path of the tendon.
Aponeurosis length was assumed to be the distance from the
tendon origin to the aponeurosis proximal end (Fig. 1).
To minimize overestimations in the displacements obtained
caused by stretch in the surrounding retinaculum upon
dorsiflexion loading (Maganaris et al., 1999), we carried out
our measurements with the ankle joint bandaged tightly with
inelastic tape. Pilot sagittal-plane magnetic resonance imaging
TA aponeurosis
TA tendon
A
A
Rest
B
P
D
.
Ankle
Knee
Fig. 1. Illustration of the displacements and lengths
measured. P and D are the proximal and distal points,
respectively, of the tibialis anterior (TA) muscle/tendon
unit. LTend and LApon are the resting lengths of the tibialis
anterior tendon and its central aponeurosis, respectively.
Arrows A and B point to the origin of the tibialis anterior
tendon and the proximal end of the central aponeurosis,
respectively. dTend and dTotal are the displacements of the
tibialis anterior tendon and tendon/aponeurosis complex,
respectively, in the transition from rest (top) to
dorsiflexion
contraction
(bottom).
Aponeurosis
displacement was estimated by subtracting dTend from
dTotal.
LTend
LApon
L
A
B
dTend
dTotal
Dorsiflexion contraction
In vivo human tendinous tissue elongation
(MRI)-based morphometrics in the ankle joint indicated that
this fixation system was effective in maintaining the resting
curved path of the tibialis anterior tendon during maximum
isometric voluntary dorsiflexion.
All measurements were taken 4 days after a familiarization
trial. All analyses were performed three times by the same
investigator, and average values were used for further analysis.
Statistical analyses
Values are presented as means ± S.D. Friedman’s rank test
was used to test differences in the tendon and aponeurosis
elongations as a function of contraction level joint moment.
The Mann–Whitney test was used to test for differences
between the tendon and aponeurosis elongations at a given
contraction level joint moment. Statistical difference was set
at a level of P<0.05.
Results
Maximal-voltage stimulation of the tibialis anterior muscle
yielded joint moment values corresponding to approximately
45 % of the moment generated upon maximum voluntary
isometric dorsiflexion. In the transition from rest to Mmax, both
the tibialis anterior tendon origin and the central aponeurosis
Fig. 2. Typical sonographs at rest (top), at 40 % of
maximal joint moment (Mmax) (middle) and at
100 % of Mmax (bottom). The white arrows point to
the tibialis anterior tendon origin (A) and the
proximal end of the tibialis anterior central
aponeurosis (B). The black arrows point to the
shadow generated by an echo-absorptive marker
attached with adhesive on the skin to confirm
the constancy of the scanning probe during
contraction.
753
proximal end shifted proximally (Fig. 2). As the dorsiflexion
joint moment increased from 20 to 100 % of Mmax, the
displacement of the tibialis anterior tendon origin at rest
increased curvilinearly from 1.3±0.4 to 4±1.3 mm (P<0.01)
(Fig. 3A). The calculated aponeurosis displacement increased
also curvilinearly from 3.7±1.3 at 20 % of Mmax to 12±4 mm
(P<0.01) at 100 % of Mmax. Tendon displacement was smaller
by between 2.4 and 8 mm (P<0.01) compared with that of the
aponeurosis at any given contraction level joint moment. The
resting lengths of the tibialis anterior tendon and central
aponeurosis were 160±16 and 170±20 mm, respectively. The
load–strain relationships of the tibialis anterior tendon and its
aponeurosis were also curvilinear (Fig. 3B). Strain increased
as a function of load from 0.8±0.1 to 2.5±0.4 % (P<0.01) in
the tendon and from 2.1±0.4 to 7±1.3 % (P<0.01) in the
aponeurosis. Tendon strain was smaller by between 61 and
64 % (P<0.01) compared with that of the aponeurosis at any
given contraction level joint moment (Fig. 3).
Reproducibility of measurements
The displacements of the tibialis anterior tendon origin and
aponeurosis proximal end in the transition from rest to 40 % of
Mmax were recorded and digitized in one subject on 10
occasions after the tendon had been preconditioned (see
754
C. N. MAGANARIS AND J. P. PAUL
Displacement (mm)
*
A
16
*
*
12
*
8
*
4
0
B
9
Strain (%)
*
*
*
6
*
*
3
0
0
20
40
60
80
100
Load (%M max)
Fig. 3. Load–displacement (A) and load/strain (B) curves of the
tibialis anterior tendinous components. Values are means ± S.D.
(N=5). An asterisk denotes a significant difference (P<0.01) between
tendon (ⵧ) and aponeurosis (䊉) elongations at a given load. Mmax,
maximal joint moment.
above). The coefficients of variation for the repeat scans were
3.3 and 5.2 %, respectively. Four different observers recorded
and analyzed the displacements of the tibialis anterior tendon
origin and aponeurosis proximal end in the transition from rest
to 20 % of Mmax in a given subject, and the coefficients of
variation were 6.1 and 8.4 %, respectively. Measurements of
the resting lengths of the tibialis anterior tendon and
aponeurosis were made 15 times and resulted in coefficients of
variation of 3.7 and 1.9 %, respectively.
Discussion
In the present study, we have demonstrated (i) that, for given
loads, the intact human tibialis anterior tendon lengthened less
than its central aponeurosis and (ii) that the tibialis anterior
tendon and aponeurosis load–elongation relationships in vivo
were curvilinear up to Mmax. Both these findings are in line
with previous reports from in vitro tensile testing of isolated
material (Butler et al., 1978; Bennett et al., 1986; Huijing and
Ettema, 1988/1989; Ettema and Huijing, 1989; Lieber et al.,
1991).
The curvilinear increase in tendinous tissue elongation at
low tension levels has been attributed to the initial crimp of
collagen fibrils (Viidik, 1973). At forces over the so-called
‘toe’ region, the stretch imposed straightens out the collagen
fibrils giving rise to linear increases in length. In vitro tendon
tests have indicated that, beyond approximately 4 % strain,
length increases in direct proportion to the applied tension and
stiffness remains unchanged until failure occurs (Butler et al.,
1978; Bennett et al., 1986). Every-day life and sporting
activities involve eccentric contractions of the tibialis anterior
muscle, but these are of low intensity and would generate
smaller forces than those elicited upon maximal isometric
contraction (Ericson et al., 1986; Jonhagen et al., 1996). In the
present experiment, the tibialis anterior tendon and aponeurosis
load–elongation curves were curvilinear, indicating that Mmax
did not exceed the tension required to reach the ‘linear’ region.
This suggests that under physiological loading the tibialis
anterior tendinous structures operate within the elastic ‘toe’
region and are not in danger of fracture.
In our experiment, the tibialis anterior tendon exhibited a
strain of 2.5 % at Mmax, while its central aponeurosis
experienced a strain of 7 %, almost three times as much as that
of the tendon. Several authors have reported values similar to
our strain values for animal and human tendons and
aponeuroses at loads equivalent to maximal muscle force (Ker
et al., 1988; Huijing and Ettema, 1988/1989; Ettema and
Huijing, 1989; Zajac, 1989; Lieber et al., 1991; Loren and
Lieber, 1995). The discrepancy in strain between the tendon
and aponeurosis is consistent across most of the studies and
indicates a decreased stiffness in the aponeurosis compared
with the tendon. Experimental evidence indicates that
differences in specimen cross-sectional area rather than in
intrinsic material compliance are probably the cause of the
strain difference between the tendon and aponeurosis (Rack
and Westbury, 1984; Scott and Loeb, 1995). However, some
studies have reported similar strain values for the tendon and
aponeurosis (Trestik and Lieber, 1993). Inter-study differences
in the methodologies followed may account for the
disagreement in the results obtained.
In our study, the aponeurosis was examined under
physiological conditions, attached to the muscle (see also
Zuurbier and Huijing, 1992; van Donkelaar et al., 1999), and
it lengthened upon muscle contraction. This finding has
important implications for muscle modelling applications. In
situ and in vivo experiments have demonstrated differences
between fibre and muscle length decreases during contraction
(Woittiez et al., 1983; Griffiths, 1991). Pennate muscle models
incorporating inextensible aponeuroses have often failed to
predict accurately the actual behaviour of muscle fibres during
contraction, thus introducing errors in the analysis of forces
(Huijing and Woittiez, 1985; Maganaris et al., 1998). The
aponeurosis compliance may partly account for the
discrepancy between modelled and experimentally measured
changes in muscle architecture. Thus, aponeurosis compliance
and also regional differences in strain along its length (see
Zuurbier et al., 1994) should be modelled accordingly.
By dividing the joint moments recorded by an average value
for the tibialis anterior moment arm of 35 mm (Maganaris et
al., 1999), we calculated that the tibialis anterior tendon tension
would increase from 70 to 550 N in the transition from 20 to
In vivo human tendinous tissue elongation
100 % of Mmax. From the slope of the tendon tension–
displacement curve, we calculated that tendon stiffness would
increase from 60 to 160 N mm−1. Aponeurosis tension and
stiffness, however, cannot be estimated from our data. In the
model shown in Fig. 1, the contractile force by which the
aponeurosis proximal end is pulled is smaller than the force
exerted on the tendon origin. It must be realized, however, that
such muscle models are oversimplified. They assume (i) a
straight-line orientation between the tendon and the
aponeurosis and (ii) that all muscle fibres are in a perfectly
parallel arrangement with no transverse inter-fibre connections
that would result in lateral force transmission. Experimental
evidence has shown that both these assumptions are invalid
(van Leeuven and Spoor, 1992; Trotter, 1993), indicating that
erroneous force estimations could be derived using simple
geometric models.
We observed large variations in the elongations of the
tendon and aponeurosis in our subjects (see Fig. 3). On
average, strain values for a given load were larger by
approximately 35 % in sedentary subjects (N=3) than in
subjects undergoing physical training (N=2). Although our
small sample size does not allow any conclusions to be drawn,
the observation of stiffer tendinous structures in response to
exercise training is consistent with previous reports from
animal studies (for a review, see Tipton et al., 1986). This
indicates that the present methodology may be sensitive
enough to detect exercise- and disuse-induced changes in intact
human tendinous structure elasticity, possibilities that are
currently being investigated systematically.
References
Bennett, M. B., Ker, R. F., Dimery, N. J. and Alexander, R. McN.
(1986). Mechanical properties of various mammalian tendons. J.
Zool., Lond. A 209, 537–548.
Butler, D. L., Grood, E. S., Noyes, F. R. and Zernicke, R. F. (1978).
Biomechanics of ligaments and tendons. In Exercise and Sport
Science Reviews, vol. 2 (ed. R. S. Hutton), pp. 125–181.
Philadelphia: The Franklin Institute Press.
Cook, C. S. and McDonagh, M. J. N. (1996). Measurement of
muscle and tendon stiffness in man. Eur. J. Appl. Physiol. 72,
380–382.
Ericson, M. O., Nisell, R. and Ekholm, J. (1986). Quantified
electromyography of lower-limb muscles during level walking.
Scand. J. Rehab. Med. 18, 159–163.
Ettema, G. J. C. and Huijing, P. A. (1989). Properties of
the tendinous structures and series elastic component of the
EDL muscle–tendon complex of the rat. J. Biomech. 22,
1209–1215.
Fukunaga, T., Ito, M., Ichinose, Y., Kuno, S., Kawakami, Y. and
Fukashiro, S. (1996). Tendinous movement of a human muscle
during voluntary contractions determined by real-time
ultrasonography. J. Appl. Physiol. 81, 1430–1433.
Griffiths, R. I. (1991). Shortening of muscle fibres during stretch of
the active medial gastrocnemius muscle: the role of tendon
compliance. J. Physiol., Lond. 436, 219–236.
Huijing, P. A. and Ettema, G. J. C. (1988/1989). Length–force
characteristics of aponeurosis in passive muscle and during
755
isometric and slow dynamic contractions of rat gastrocnemius
muscle. Acta Morphol. Neerl.-Scand. 26, 51–62.
Huijing, P. A. and Woittiez, R. D. (1985). Notes on planimetric and
three-dimensional muscle models. Neth. J. Zool. 35, 521–525.
Ito, M., Kawakami, Y., Ichinose, Y., Fukashiro, S. and Fukunaga,
T. (1998). Nonisometric behavior of fascicles during isometric
contractions of a human muscle. J. Appl. Physiol. 85, 1230–1235.
Jonhagen, S., Ericson, M. O., Nemeth, G. and Erikkson, E. (1996).
Amplitude and timing of electromyographic activity during
sprinting. Scand. J. Med. Sci. Sports 6, 15–21.
Kawakami, Y., Abe, T. and Fukunaga, T. (1993). Muscle fiber
pennation angles are greater in hypertrophied than in normal
muscles. J. Appl. Physiol. 74, 2740–2744.
Ker, R. F., Alexander, R. McN. and Bennett, M. B. (1988). Why
are mammalian tendons so thick? J. Zool., Lond. 216, 309–324.
Lieber, R. L., Leonard, M. E., Brown, C. G. and Trestik, C. L.
(1991). Frog semitendinosis tendon load–strain and stress–strain
properties during passive loading. Am. J. Physiol. 30, C86–C92.
Loren, G. J. and Lieber, R. L. (1995). Tendon biomechanical
properties enhance human wrist muscle specialization. J. Biomech.
28, 791–799.
Maganaris, C. N., Baltzopoulos, V. and Sargeant, A. J. (1998). In
vivo measurements of the triceps surae architecture in man:
implications for muscle function. J. Physiol., Lond. 512, 604–613.
Maganaris, C. N., Baltzopoulos, V. and Sargeant, A. J. (1999).
Changes in the tibialis anterior tendon moment arm from rest to
maximum isometric dorsiflexion: In vivo observations in man. Clin.
Biomech. 14, 661–666.
Morgan, D. L. (1977). Separation of active and passive components:
Short-range stiffness of muscle. Am. J. Physiol. 232, C45–C49.
Narici, M. V., Binzoni, T., Hiltbrand, E., Fasel, J., Terrier, F. and
Cerretelli, P. (1996). In vivo human gastrocnemius architecture
with changing joint angle at rest and during graded isometric
contraction. J. Physiol., Lond. 496, 287–297.
Rack, P. M. H. and Westbury, D. R. (1984). Elastic properties of
the cat soleus tendon and its effects on mechanical properties. J.
Physiol., Lond. 282, 253–261.
Scott, S. H. and Loeb, G. E. (1995). Mechanical properties of
aponeurosis and tendon of the cat soleus muscle during wholemuscle isometric contractions. J. Morph. 224, 73–86.
Shorten, M. R. (1987). Muscle elasticity and human performance. In
Current Research in Sport Sciences. Medicine and Sports Science,
vol. 25 (ed. B. van Gheluwe and J. Atha), pp. 1–18. Basel: Karger.
Siegler, S., Moskowitz, G. D. and Freedman, W. (1984). Passive
and active components of the internal moment developed about the
ankle joint during ambulation. J. Biomech. 19, 647–652.
Smith, C. V., Young, I. S. and Kearney, J. N. (1996). Mechanical
properties of tendons: Changes with sterilization and preservation.
J. Biomech. Eng. 118, 56–61.
Tipton, C. M., Vailas, A. C. and Matthes, R. D. (1986).
Experimental studies on the influence of physical activity on
ligaments, tendons and joints: A brief review. Acta Med. Scand.
Suppl. 71, 157–168.
Trestik, C. L. and Lieber, R. L. (1993). Relationship between
Achilles tendon mechanical properties and gastrocnemius muscle
function. J. Biomech. Eng. 115, 225–230.
Trotter, J. A. (1993). Functional morphology of force transmission
in skeletal muscle. Acta Anat. 146, 205–222.
van Donkelaar, C. C., Willems, P. J. B., Muijtjens, A. M. M. and
Drost, M. R. (1999). Skeletal muscle transverse strain during
isometric contraction at different lengths. J. Biomech. 32, 755–762.
756
C. N. MAGANARIS AND J. P. PAUL
van Leeuwen, J. L. and Spoor, C. W. (1992). Modelling
mechanically stable muscle architectures. Phil. Trans. R. Soc. Lond.
B 336, 275–292.
Viidik, A. (1973). Functional properties of collagenous tissues. Int.
Rev. Connective Tissue Res. 6, 127–215.
Woittiez, R. D., Huijing, P. A. and Rozendal, R. H. (1983).
Influence of muscle architecture on the length–force diagram of
mammalian muscle. Pflügers Arch. 399, 275–279.
Zajac, E. F. (1989). Muscle and tendon: Properties, models, scaling
and application to biomechanics and motor control. CRC Crit. Rev.
Biomed. Eng. 17, 359–411.
Zuurbier, C. J., Everard, A. J., Van der Wees, P. and Huijing, P.
A. (1994). Length–force characteristics of the aponeurosis in the
passive and active muscle condition and in the isolated condition.
J. Biomech. 27, 445–453.
Zuurbier, C. J. and Huijing, P. A. (1992). Influence of muscle
geometry on shortening speed of fibre, aponeurosis and muscle. J.
Biomech. 25, 1017–1026.