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First published online July 20, 2007
Journal of Experimental Biology 210, 2743-2753 (2007)
Published by The Company of Biologists 2007
doi: 10.1242/jeb.003814
Adaptational responses of the human Achilles tendon by modulation of the applied cyclic strain magnitude
German Sport University of Cologne, Institute of Biomechanics and Orthopaedics, Carl-Diem-Weg 6, 50933 Cologne, Germany
* Author for correspondence (e-mail: Arampatzis{at}dshs-koeln.de)
Accepted 14 May 2007
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Summary |
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 TOP Summary IntroductionMaterials and methodsResultsDiscussionReferences |
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Key words: MRI, ultrasonography, tendon plasticity, in vivo, exercise, strain
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Introduction |
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TOPSummaryIntroductionMaterials and methodsResultsDiscussionReferences |
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;
Hsieh et al., 2000;
Zeichen et al., 2000). It has
been reported that the cells sense the applied strain
(Chiquet, 1999;
Chiquet et al., 2003) and
regulate the synthesis of matrix proteins
(Arnoczky et al., 2002;
Skutek et al., 2003;
Miller et al., 2005), the gene
expression of proteoglycans and collagen
(Robbins and Vogel, 1994;
Kim et al., 2002;
Hsieh et al., 2000), the
alignment and density of the collagenous matrix
(Pins et al., 1997;
Wang et al., 2001;
Wang et al., 2003a) as well as
the expression of several growth factors
(Skutek et al., 2001;
Yang et al., 2004;
Olesen et al., 2006). Studies
examining the influence of mechanical loading on the regulation of the
biosynthesis of connective soft tissues demonstrated that the strain
magnitude, strain frequency, strain rate and strain duration of cells
influence the cellular biochemical responses
(Skutek et al., 2003;
Arnoczky et al., 2002;
Yang et al., 2004) and the
mechanical properties of collagen fascicles
(Yamamoto et al., 2002;
Yamamoto et al., 2003;
Yamamoto et al., 2005).
Furthermore, the above studies illustrate the highly plastic nature of fibrous
connective tissues and provide evidence that strain of tendon cells is an
important regulator for the homeostasis of connective tissues.
More than two decades ago, Woo et al.
(Woo et al., 1982) formulated
the hypothesis that the homeostatic responses of soft tissues subjected to
mechanical loads may be represented by a non-linear curve. Immobilisation
causes a rapid decline in the mechanical properties whereas long-term exercise
initiates a slight increase in mechanical properties compared with normal
daily activities (Woo et al.,
1982). More recently, it has been suggested that the applied
strain on the connective tissues may have a threshold or set-point to create a
homeostatic perturbation in the collagenous matrix that regulates the
catabolic and anabolic responses of the cells
(Brown et al., 1998;
Lavagnino and Arnoczky, 2005).
An external mechanical loading of the tissue above the upper limit at which
the endogenous contraction of the fibroblasts may maintain their tensional
homeostasis should stimulate cells for remodelling, whereas a reduction of the
mechanical loading below the lower limit will lead to tissue destruction
(Lavagnino and Arnoczky, 2005;
Lavagnino et al., 2006).
In agreement with both the hypothesis formulated by Woo et al.
(Woo et al., 1982) and the
concept of the `homeostatic calibration point'
(Lavagnino and Arnoczky, 2005;
Lavagnino et al., 2006), which
supports the existence of an upper and a lower limit determining a homeostatic
perturbation, we did not find a graded response between exercise intensity and
mechanical properties of the human triceps surae tendon and aponeurosis by
comparing sprinters, endurance runners and subjects not being active in sports
(Arampatzis et al., 2007a).
Only the sprinters group showed a higher stiffness at the triceps surae tendon
and aponeurosis compared with the other two groups examined
(Arampatzis et al., 2007a).
Further, we suggested (Arampatzis et al.,
2007a) that the mechanical properties of the human triceps surae
tendon and aponeurosis remain at control level in a wide range of applied
strains and that the strain magnitude, strain frequency and strain rate should
exceed a given threshold in order to trigger additional adaptation effects.
Although numerous important studies have previously demonstrated the
plasticity of human tendons in response to resistance exercise
(Kubo et al., 2001a;
Kubo et al., 2001b;
Kubo et al., 2002;
Reeves et al., 2003a;
Reeves et al., 2003b;
Reeves et al., 2005), and even
though it is well accepted that tendons are able to remodel their mechanical
and morphological properties in response to mechanical loading, there is
little information about the effects of controlled modulation in cyclic strain
magnitude, frequency or rate applied to the tendon on the adaptation of the
mechanical and morphological properties of tendons in vivo. Thus, it
may be concluded that the tendon responses to different cyclic strain
magnitudes in vivo remain a fundamental unanswered question.
Knowledge of tendon plasticity in response to the magnitude of the mechanical
load induced as cyclic strain applied to the tendon may help improve the
intervention process of tendon adaptation and tendon healing.
The purpose of this study was to examine the effect of two different
exercise interventions of cyclic strain applied to the Achilles tendon on the
adaptation of its mechanical and morphological properties. Both interventions
were performed at the same frequency and volume but at different magnitudes of
tendon strain (2.85±0.99% vs 4.55±1.38% strain). Based
on reports about human tendon plasticity
(Kubo et al., 2001a;
Kubo et al., 2001b;
Kubo et al., 2002;
Reeves et al., 2003a;
Reeves et al., 2003b;
Reeves et al., 2005), the
concept of the homeostatic calibration point
(Lavagnino and Arnoczky, 2005;
Lavagnino et al., 2006) and
the non-graded response of the mechanical properties of the human triceps
surae tendon and aponeurosis in an intensity-dependent manner of sport
activity (Arampatzis et al.,
2007a), we expected an adaptation effect on the Achilles tendon
only after the high-strain-magnitude exercise intervention, demonstrating a
threshold in strain magnitude for further adaptational effects in
vivo.
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Materials and methods |
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Exercise protocol
The intervention lasted 14 weeks. Four times per week the experimental
group performed five sets of repetitive (3 s loading, 3 s relaxation),
isometric plantar flexion contractions (ankle angle at 85° dorsal flexion,
knee angle fully extended at 180° and the hip flexed at 140°).
Repetitive isometric plantar flexion contractions were used to induce cyclic
strains on the triceps surae tendon and aponeurosis. The participants
exercised one leg at low-magnitude tendon–aponeurosis strain
(low-strain-magnitude exercise) and the other leg at high-magnitude
tendon–aponeurosis strain (high-strain-magnitude exercise). The
assignment of low and high strain exercise to each leg was random. Based on
earlier experience (Arampatzis et al.,
2005a; Mademli et al.,
2006), we predicted that a plantar flexion moment at 55% of the
achieved maximum moment during a maximum voluntary contraction (MVC) should
induce a tendon–aponeurosis strain between 2.5 and 3.0% whereas a
plantar flexion moment at 90% of the MVC should induce between 4.5 and 5.0%
strain. At each set, the leg exercised at high strain magnitude performed four
repetitions (3 s loading, 3 s relaxation) at 90% of the MVC whereas the other
leg (low strain magnitude) performed seven repetitions at 55% of the MVC
(Fig. 1). This way (four
vs seven repetitions per set for the high- and the
low-strain-magnitude exercise, respectively), both legs were trained at the
same exercise volume (integral of the plantar flexion moment over time). The
above experimental design provided an intervention of similar frequency and
volume but different magnitude of applied cyclic strain to the triceps surae
tendon and aponeurosis of each trained leg.
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Measurement of plantar flexion moment and voluntary activation
The subjects were seated on a dynamometer (Biodex-System3) with the ankle
angle in a dorsal flexed position at 85° (tibia perpendicular to the sole,
corresponding to 90° ankle angle), the knee fully extended at 180° and
the hip flexed at 140°. In this position, the subjects performed maximal
isometric plantar flexion contractions. After a warm-up period, consisting of
2–3 min submaximal isometric contractions, and three MVCs the
participants were instructed to produce a maximal isometric force ramp with
the highest possible rate of force generation. We used the three MVCs to
exclude the preconditioning effect on the tendon strain–force
relationship. The three MVCs at the beginning of the intervention should not
have any substantial training effect. The twitch interpolation technique
(Merton, 1954) was used to
determine the VA of the plantar flexor muscles during the contraction. We
evoked a superimposed twitch (three 500 µs square-wave pulses separated by
5 ms) at the plateau of the MVC and three supramaximal twitches after the MVC
when the plantar flexor muscles were relaxed
(Fig. 2) using a stimulator
(Model DS7A digitimer; Digitimer Ltd, Welwyn Garden City, Hertfordshire, UK).
The voluntary activation was calculated by normalising the evoked interpolated
twitch torque (ITT) to the mean of the three resting twitch torques (RTT):
VA=[1–ITT/RTT)x100].
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The resultant moments at the ankle joint were calculated through inverse
dynamics. The method for calculating the resultant joint moments has been
previously described (Arampatzis et al.,
2005b). Kinematic data were recorded using the Vicon 624 system
(Vicon Motion Systems, Oxford, United Kingdom) with eight cameras operating at
120 Hz. To calculate the lever arm of the ankle joint during ankle plantar
flexion, the centre of pressure under the foot was determined by means of a
flexible pressure distribution insole (Pedar-System, Novel GmbH, Munich,
Germany) operating at 99 Hz (Arampatzis et
al., 2005b). The compensation of moments due to gravitational
forces was determined for all subjects before each ankle plantar flexion
contraction. The antagonistic moment of the tibialis anterior (TA) during MVC
was estimated by establishing a relationship between electromyographic (EMG)
activity and exerted moment for the TA, while working as agonist
(Mademli et al., 2004). This
was established by measuring EMG and moment during relaxation and during two
submaximal ankle dorsiflexion contractions
(Mademli et al., 2004).
Measurement of tendinous tissue elongation
After the MVC with the superimposed twitches, the participants were
instructed to produce another maximal isometric force ramp, gradually
increasing the plantar flexion effort over 3 s (loading), and to hold the
achieved moment for about 2–3 s. A 7.5 MHz linear array ultrasound probe
(Aloka SSD 4000; Tokyo, Japan; 43 Hz) was used to visualise the distal tendon
and aponeurosis of the gastrocnemius medialis (GM) during the MVC. The
ultrasound images were recorded on video tapes for further analysis. On the
video images, a clear visible cross-point (intersecting point between the
distal aponeurosis and a fascicle of the GM muscle) was identified and its
displacement was measured in relation to a skin marker
(Fig. 3). The exact protocol
for analysing the tendinous tissue elongation during ankle plantar flexion is
described in detail elsewhere (Arampatzis
et al., 2005a). Briefly, the ultrasound probe was placed above the
muscle belly at about 50% of its length. For the analysis of the video tapes
every single frame was digitised using video analysis software (Simi Motion
5.0; SIMI Reality Motion System GmbH, Unterschleißheim, Germany). The
effect of inevitable joint angular displacement on the observed elongation of
the tendon and aponeurosis during the MVC was taken into account by capturing
the motion of the tendons and aponeuroses from the GM during a passive
(inactive) motion of the ankle joint
(Muramatsu et al., 2001). The
passive motion of the ankle joint has been analysed during the plantar flexion
because the angular rotation at the ankle joint during the `isometric' MVC was
also a plantar flexion. The error of this method on the strain value is
0.3% and, thus, has a negligible effect on the examined in vivo
strain of the tendon and aponeurosis
(Arampatzis et al., 2007b). The
analysed cross-point at the aponeurosis was digitised during the inactive
condition at the same ankle angle changes as observed during the MVC
(Arampatzis et al., 2005a). The
elongation of the GM tendon and aponeurosis was calculated as the difference
between the measured and the passive (due to joint rotation) displacement of
the analysed point at the aponeurosis (Fig.
4).
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In order to estimate the resting length of the GM tendon and aponeurosis,
the subjects were seated on the dynamometer with the knee at 180° and the
ankle at 110°. We used this specific position because De Monte et al.
(De Monte et al., 2006)
reported the existence of slackness in the inactive GM muscle–tendon
unit between 121° and 107° ankle angle and 180° knee angle and
that the 110° ankle angle is a suitable position to examine the resting
length of the GM tendon and aponeurosis. The length of the curved path from
the tuberositas calcanei (defined as the origin of the Achilles tendon) to the
skin marker (Fig. 3) was
measured along the skin using flexible measuring tape. Thus, the resting
length of the GM tendon and aponeurosis was defined as the length of the path
between the tuberositas calcanei and the analysed cross-points identified on
the ultrasound images. The tendon force was calculated by dividing the plantar
flexion moment by the tendon moment arm. The moment arm of the Achilles tendon
was calculated using the data provided by Maganaris et al.
(Maganaris et al., 1998). The
elongation and strain of the tendon and aponeurosis during the MVC was
identified and analysed at the maximum calculated tendon force and at every
100 N. The stiffness of the triceps surae tendon and aponeurosis has been
calculated as the slope of the calculated tendon force vs
tendon–aponeurosis elongation between 50% and 100% of the maximum tendon
force by means of linear regressions.
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To standardise the levels of the transversal images, two landmarks, the most proximal aspect of the tuberositas calcanei and the most distal aspect of the soleus muscle, were utilised. The sagittal images served to obtain the locations of both points. On each transversal image, the boundaries of the Achilles tendon were outlined manually using the software 3D Doctor (Able Software Corp., Lexington, MA, USA). The tendon boundaries and the coordinates of the two landmarks were exported and processed using Matlab (The Mathworks, Natick, MA, USA). For each of the subsequent cross-sections, the area and the location of the centroid were calculated. The length of the Achilles tendon was calculated as the curved path through the centroids of the cross-sections between the two landmarks. The CSA of the Achilles tendon was identified and analysed at every 10% of tendon length. To examine the elastic modulus of the Achilles tendon we calculated the relationship between tendon stress and tendon–aponeurosis strain from 50% to 100% of the maximum tendon stress by means of linear regressions. To calculate the tendon stress (tendon force/tendon CSA) we used the average value of the CSA of the Achilles tendon from 10% to 100% of the tendon length.
Statistics
A T-test for two dependent samples was used to check the
intervention-related differences in the examined parameters (maximum plantar
flexion moment, voluntary activation, tendon–aponeurosis strain at every
100 N and CSA of the Achilles tendon at every 10% of the tendon length) in
each group. Further, to check the ratios (post- to pre-exercise values) of the
examined parameters we used a one-way analysis of variance (ANOVA) and
Bonferroni post-hoc comparisons between the three groups (control
group, no exercise intervention; experimental group 1, low-strain-magnitude
exercise intervention; experimental group 2, high-strain-magnitude exercise
intervention). The level of significance for all comparisons was set to
=0.05. In all figures, the data are presented as means ±
standard error of mean (s.e.m.), whereas in the text and tables they are
expressed as means ± standard deviation (s.d.).
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After the 14 weeks intervention applying cyclic loading to the Achilles tendon, the tendon–aponeurosis strain for a given tendon force (every 100 N) did not show any statistically significant (P>0.05) changes in the low-strain-exercised leg (Fig. 7), indicating no alteration in the strain–force relationship of the tendon and aponeurosis due to the intervention. By contrast, after the 14 weeks intervention at high strain magnitude, the strain values for a given tendon force (every 100 N) up to 600 N were significantly (P<0.05) lower than before (Fig. 7), demonstrating a higher gradient in the force–strain curve as compared with the pre-exercise curve. As expected, the control group did not show any differences in the strain–force relationship before and after the 14-week period (Fig. 7). Up to 1200 N tendon force, the ratios of the strain values (post- to pre-exercise) were significantly (P<0.05) lower for the leg loaded with the high strain magnitude than for the low-strain-exercised leg and the control group (Fig. 7).
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;
Wang and Thampatty, 2006;
Lavagnino et al., 2006).
In the literature, it is well accepted that mechanical load induced as
cyclic strain on connective soft tissues such as tendons is an important
regulator of the fibroblast's metabolic activity as well as a regulator of the
maintenance of the tendon matrix (Arnoczky
et al., 2002; Barkhausen et
al., 2003; Screen et al.,
2005; Webb et al.,
2006). Furthermore, the modulation of the mechanical stimuli
affects several physiological parameters of human fibroblasts and coordinates
the amount of proliferation, apoptosis and expression of proteins
(Zeichen et al., 2000;
Skutek et al., 2001;
Skutek et al., 2003;
Barkhausen et al., 2003). For
example, loading of tendon cells causes a downregulation of catabolic gene
expression and an upregulation of anabolic gene expression
(Lavagnino and Arnoczky, 2005;
Lavagnino et al., 2006)
whereas immobilisation promotes catabolic responses (i.e. degeneration of the
extracellular matrix) imposed by an upregulation of matrix metalloproteinases
(Amiel et al., 1982;
Hannafin et al., 1995;
Brown et al., 1998;
Arnoczky et al., 2004).
Although there is little information in the literature about the effects of
controlled tendon strain magnitudes on the homeostatic perturbation and the
induced adaptational responses of tendons in vivo, in vitro studies
have demonstrated the existence of a threshold in tendon strain magnitude for
triggering fibroblast proliferation (Yang
et al., 2004), stimulation of the gene expression of inflammatory
mediators such as interstitial collagenase
(Lavagnino et al., 2003) or
prostaglandin E2 (Wang et al.,
2003b) and changes in the elastic modulus and tensile strength of
cultured collagen fascicles after loading
(Yamamoto et al., 2003).
Recently, Kubo et al. (Kubo et al.,
2006) reported an increase in human vastus lateralis
tendon–aponeurosis stiffness after high-load isokinetic training of the
knee extensor muscles (80% of the isokinetic MVC) but no changes in
tendon–aponeurosis stiffness after low-load isokinetic knee extension
training (20% of the isokinetic MVC), which is in agreement with our
results.
The homeostatic perturbation in the connective tissues induced by
mechanical loading affects several biochemical cellular responses
(Robbins and Vogel, 1994;
Hsieh et al., 2000;
Kim et al., 2002). The concept
of homeostatic calibration point
(Lavagnino and Arnoczky, 2005;
Lavagnino et al., 2006)
predicts that mechanical loading of the tendon above the homeostatic
calibration point (upper limit) will trigger anabolic responses whereas a
reduction of tendon loading below the homeostatic level (lower limit) will
lead to catabolic cell responses. The findings of the present study showing
adaptational effects on the Achilles tendon only at the leg exercised at a
high strain magnitude indicate that the mechanical load applied to the leg
exercised at a low tendon strain magnitude did not influence the existing
internal tensional homeostasis of the tendon cells regulating the anabolic or
catabolic responses. The results further show that the mechanical load exerted
on the Achilles tendon during the low-strain-magnitude exercise is no more a
sufficient stimulus for triggering further adaptation effects on the Achilles
tendon than the stimulus provided by the mechanical load applied during daily
activities. Furthermore, our findings indicate that the 4.55% strain applied
during the high-strain-magnitude intervention was above the homeostatic
calibration point and thus was sufficient to elicit a homeostatic perturbation
at the Achilles tendon that triggered anabolic cell responses, causing the
changes observed at the tendon–aponeurosis strain–force
relationship and the region-specific hypertrophy of the tendon. In the present
study, we controlled the strain magnitude, strain frequency and the exercise
volume but not the strain rate during the interventions. The participants
achieved the target moment as fast as possible and, therefore, the strain rate
should not be very different between the two examined interventions. However,
based on our experimental design it is not possible to investigate a potential
effect of the strain rate on the tendon adaptational responses that we
discovered.
In the present study, we found that 14 weeks exercise at high strain
magnitude had a clear influence on the strain–force relationship of the
tendon–aponeurosis unit and led to an increase of the CSA of the
Achilles tendon at 60 and 70% of its length. The region-specific hypertrophy
of the Achilles tendon may partly explain the changes in the
strain–force relationship of the tendon–aponeurosis unit but not
the increase in tendon elastic modulus after the intervention. Besides tendon
hypertrophy, there are some other adaptation possibilities that may affect the
tendon stress–strain relationship. The organisation of the extracellular
matrix components includes mechanisms transmitting tensile forces along the
interfibrilar matrix. Several studies have reported that cells have the
ability to produce a better organised collagen matrix modulated by cyclic load
(Steinmeyer and Knue, 1997;
Wang and Grood, 2000;
Wang et al., 2003a;
Webb et al., 2006) and this
way achieve an increase in tissue stiffness
(Brown et al., 1998;
Lo et al., 2000). The methods
used in the present study do not permit examination of such adaptation
possibilities at the Achilles tendon; nevertheless, the clear changes in the
tendon–aponeurosis strain–force relationship, the only
region-specific tendon hypertrophy, the increase of the tendon elastic
modulus, as well as reports of other studies demonstrating an increase in
human tendon stiffness and elastic modulus with no changes in the tendon's CSA
(Kubo et al., 2002;
Reeves et al., 2003a) provide
evidence for the plasticity of the organisation of the tendon's extracellular
matrix in vivo (i.e. density of matrix proteins, cell orientation,
proteoglycan content and composition).
The maximum plantar flexion moment increased after the intervention in both
exercised legs (on average 20 and 33% for the low- and high-strain-magnitude
exercised legs, respectively). This is in agreement with other studies
reporting an increase in muscle strength after low- and high-intensity
resistance training (Kaneko et al.,
1983; Takarada et al.,
2000; Moore et al.,
2004). The voluntary activation of the plantar flexor muscles
during the MVC were quite high (97–99%) and did not show any differences
before and after both exercise interventions. This indicates that the increase
in muscle strength observed after the 14-week intervention was not due to
neuronal factors. In the leg exercised at high strain magnitude the maximum
tendon–aponeurosis strain during the MVC did not differ before and after
the intervention. These findings, namely an increase in muscle strength with
no changes in maximum tendon–aponeurosis strain, suggest a coordinated
muscle–tendon unit adaptation at the high-strain-magnitude intervention.
Recently, Miller et al. (Miller et al.,
2005) reported similar changes in the time course of tendon
collagen and myofibrillar protein synthesis rates after non-damaging exercise,
supporting a coordinated musculotendinous adaptation. However, the results of
the leg exercised at low strain magnitude (i.e. increase in muscle strength
with no changes in the tendon–aponeurosis strain–force
relationship) do not show any coordinated muscle–tendon unit adaptation.
This indicates that the threshold of mechanical loading necessary to trigger
adaptational effects is higher for the tendon than for the muscle.
We found an increase in tendon–aponeurosis stiffness, an increase in
tendon elastic modulus and a region-specific hypertrophy of the Achilles
tendon after the high-strain-magnitude exercise (i.e. 90% MVC). The reported
maximal plantar flexion moment values calculated by inverse dynamic during
daily activities such as walking are about 120–130 Nm
(Winter, 1984). This means
that the resultant maximal ankle plantar flexion joint moment while walking is
similar or even higher than the applied plantar flexion moment at the
high-strain-magnitude intervention. Therefore, it can be argued that the
mechanical load on the Achilles tendon at the high-strain-magnitude
intervention was not higher compared with normal walking and, thus, the
applied mechanical stimulus does not explain the adaptational effects.
However, it is difficult to compare the loading on the Achilles tendon induced
by walking with the loading induced during the examined strength training. The
maximal ankle plantar flexion joint moments while walking do not last for long
(instantaneous values). The mean ankle plantar flexion joint moments while
walking are about 50–55 Nm
(Karamanidis and Arampatzis,
2007). The duration of the loading is also different between
walking (600 ms) and the exercise intervention used in the present study
(3 s). Further, the resultant joint moments during daily activities calculated
by inverse dynamics are not only compensated by muscles. Passive structures as
well as contact forces between bones also absorb parts of these moments. For
example, the maximal ankle plantar flexion joint moment while walking occurs
at a joint position of 15–20° dorsiflexion
(Winter, 1984). At this ankle
joint angle, the passive joint moment can achieve values between 20 and 30 Nm
(Riener and Edrich, 1999;
Mullaney et al., 2006).
Moreover, due to the viscoelastic behaviour of the connective tissues, in a
dorsiflexed position the passive ankle joint moments may be higher in a
dynamic condition (Gajdosik et al.,
2005) such as walking. Studies examining the EMG activity of the
triceps surae muscles while walking reported values between 19 and 42% of the
maximal isometric EMG value (Ericson et
al., 1986). Given that the muscle force depends at least on the
force potential due to the force–length–velocity relationship and
the activation level (Winters,
1990), submaximal EMG activity suggests submaximal muscle
forces.
In conclusion, our results demonstrate a decrease in
tendon–aponeurosis strain at a given tendon force and a region-specific
hypertrophy of the Achilles tendon after 14 weeks of high-strain-magnitude
exercise (4.6% tendon–aponeurosis strain) and no changes in tendon
properties after the same period of low-strain-magnitude exercise (2.9%
tendon–aponeurosis strain) of similar frequency and volume. The
contractile capacity of the plantar flexor muscles increased in both levels of
exercised legs but the increase was higher at the high-strain-magnitude than
at the low-strain-magnitude exercise. The results further show that the strain
magnitude applied to the human Achilles tendon should exceed a given threshold
to trigger adaptational effects on the mechanical and morphological properties
of the tendon and that applied strains with low magnitude (2.5–3.0%) are
not a sufficient stimulus to trigger adaptation effects on the Achilles tendon
beyond those triggered by the mechanical load applied during daily
activities.
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Acknowledgments |
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References |
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TOPSummaryIntroductionMaterials and methodsResultsDiscussionReferences |
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Statistically
significant differences between low- and high-strain-magnitude exercised legs
(P<0.05).
