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The Journals of Gerontology Series A: Biological Sciences and Medical Sciences 59:B441-B448 (2004)
© 2004 The Gerontological Society of America

Gender Differences in Human Muscle and Joint Mechanical Properties During Plantar Flexion in Old Age

Julien Ochala1,, Denis Valour1, Michel Pousson1, Daniel Lambertz2 and Jacques Van Hoecke1

1 Equipe Motricité-Plasticité, Faculté des Sciences du Sport, Université de Bourgogne, Dijon, France.
2 Département de Génie Biologique, Université de Technologie, Compiègne, France.

Address correspondence to Julien Ochala, Equipe INSERM-ERM 207 Motricité-Plasticité, Faculté des Sciences du Sport, Université de Bourgogne, BP 27 877, 21 078 Dijon Cedex, France. E-mail: julien.ochala{at}u-bourgogne.fr


   Abstract
Top
 Abstract
Methods
 Results
 Discussion
 References
 
Muscle and joint mechanical properties during plantar flexion were investigated in 13 elderly women (EW) (age range 73–83 years) and 15 elderly men (EM) (age range 74–81 years). Maximal torque, at several angular velocities, was measured to construct torque–velocity relationship. This led to the calculation of an index of maximal shortening velocity (VImax) at 10% of maximal isometric torque. Two methods were then used to calculate musculotendinous (quick-release movements) and musculoarticular (sinusoidal perturbations) stiffness. In both cases, stiffness was linearly related to torque, and the slope was defined as a stiffness index (SI): SIMT and SIMA, respectively. Maximal torques as well as VImax (p <.05) were lower in EW compared with EM. Furthermore, SIMT and SIMA values were higher for EW than for EM (p <.05). These results are interpreted in terms of possible differences in the muscle mass, fiber-type distribution, and tendons. They may also have implications for daily motor behavior.

THE assessment of muscle mechanical properties in elderly women (EW) and elderly men (EM) is mostly limited to an evaluation of muscle contractility. For instance, EW tend to have smaller absolute strength values than EM, i.e., lower maximal voluntary torques at the same velocity or across a range of velocities under isokinetic conditions on elbow flexors and extensors (1) as well as knee flexors and extensors (2,3). The inability for women to develop the same voluntary torques as men induces numbers of functional disorders in daily motor tasks, such as rising from a chair or walking, increasing the risk of falls and hip fractures (4). In the literature, the lower maximal voluntary torques produced by EW have been often attributed to hormone deficiency (estrogen and progestogen) at the menopause (5,6) and to gender differences in muscle mass and/or thickness of the adipose tissue layer. Aging similarly affects muscles of women and men, e.g., reduced numbers of both slow-twitch and fast-twitch muscle fibers, related to loss of {alpha}-motoneurons and incomplete reinnervation of denervated muscle cells (7). Nevertheless, some studies reported that slow-twitch fibers occupied a greater area in women, whereas fast-twitch fibers occupied a greater area in men (8,9), suggesting that gender differences may also exist at the single muscle fiber level (10).

Quantification of muscle mechanical properties requires the assessment of both muscle contractile and elastic properties. By using a classical model of muscle mechanics (11), muscle elastic properties can be described by the series elastic component (SEC), which is composed of an active fraction (i.e., muscle fibers) and a passive fraction (i.e., tendon) (12). SEC behavior is obtained by means of a quick-release technique and SEC stiffness, i.e., the ratio between changes in force and changes in length, is commonly achieved by the tension–extension relationship (e.g., 13). The heterogeneity of the SEC leads to nonlinear tension–extension relationship, and stiffness of the SEC is classically linearly related to force.

Differences in the stiffness of the SEC, i.e., musculotendinous stiffness, have already been reported in humans, after periods of changes in functional demands, as a result of hyperactivity (14), microgravity (15), and disease (16). Furthermore, differences in musculotendinous stiffness between young men and elderly men have also been reported (17). The reported changes in musculotendinous stiffness were interpreted in terms of adaptation of the properties of the muscles and tendons involved. Thus, assessing musculotendinous stiffness in EW and EM may allow for detecting gender differences in the muscles and tendons, which can induce changes in motor behavior. For instance, musculotendinous stiffness influences muscle efficiency, i.e., force transmission via tendon directly to bone, notably when eccentric muscle actions precede concentric contractions, i.e., stretch-shortening cycle (SSC) involved during most daily activities, e.g., walking (18). In fact, the transmission of force generated by the contractile elements occurs across the musculotendinous deformation (11). Higher musculotendinous stiffness may be suitable for transmitting the force more effectively (19) and may favor the release of potential energy during SSC, since it shortens the coupling time, i.e., the time delay between stretching and shortening phases (18).

Muscle and joint can also be considered as a whole and the dynamic mechanical properties are described by joint dynamics. Joint dynamics deal with the relation between the torque acting about the joint and the angular position of a joint, which represent musculoarticular stiffness. The involved structures can be the SEC and articular structures (15). Musculoarticular stiffness represents an important parameter for posture and movement control because it governs the mechanics of the interaction between the musculoskeletal system and the external environment. An approach to measure musculoarticular stiffness properties is to apply sinusoidal disturbances over a range of frequencies. This technique has been used to quantify alterations due to changes in functional demands (15,20) and disease (16) and has allowed for determining the role played by muscles, tendons, and joints in movement.

The present study was designed to determine whether, in humans, gender differences in musculotendinous and musculoarticular stiffness during plantar flexion in old age exist and whether the differences are consistent with anatomical and physiological data reported in the literature. Furthermore, force production capacities of the ankle plantar flexor muscles under isometric and isokinetic conditions were also measured.


   METHODS
Top
 Abstract
 Methods
Results
 Discussion
 References
 
Participants
The study was performed on 28 healthy elderly participants organized into 2 groups: elderly women (EW) (13 women, age range 73–83 years) and elderly men (EM) (15 men, age range 74–81 years). The male and female volunteers were sedentary and free from any neurological deficit and/or orthopedic disability. Physical characteristics of these groups are shown in Table 1. All procedures were explained to the participants verbally, and a written consent form was completed before the study procedures were begun. The protocol was approved by a local ethics committee, and the study was carried out according to the guidelines of the Declaration of Helsinki.


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  		Table 1. Physical Characteristics of Participants (Means ± SD).

 
Materials
An ankle ergometer was used for the study. It was especially designed to test the mechanical properties of plantar flexor muscles and the ankle joint (21). Briefly, the ergometer consisted of 2 main units: a) a power unit that contained the actuator, its power supply unit, position, velocity, and torque transducers, and its associated electronics; and b) a driving unit controlled by a computer equipped with a specific 12-bit A/D timer board. Angular displacement was measured with an optical digital sensor, and angular velocity was captured from a resolver bound to the rotor, except for velocities greater than 15.70 rad/s, which required a tachometer. Torque was obtained by using a strain-gauge torque transducer. Specific menu-driven software controlled all procedures, and recorded mechanical variables and electromyograms for later analysis. Mechanical variables and electromyograms (EMGs) were loaded directly onto a computer and instructions were presented to the participant on an oscilloscope.

Study Protocol
The participant lay comfortably on an adjustable table, in a supine position, with the left ankle placed at 1.56 rad, i.e., neutral position (with 0 rad being the perpendicularity of the ergometer footplate to the long axis of the tibia), and attached rigidly to the actuator of the ankle ergometer. The horizontal bimalleolar axis coincided with the axis of rotation of the actuator. The knee was flexed to 2.08 rad to minimize the contribution of the gastrocnemii muscles (22). The shoulders were maintained by special shoulder holders. Surface EMGs were recorded by using bipolar silver chloride surface electrodes of 10 mm diameter, with an interelectrode (center-to-center) distance of 20 mm. For the soleus, recording electrodes were placed along the middorsal line of the leg, about 2 cm distal from where the 2 heads of the gastrocnemii join the Achilles tendon. For the gastrocnemii and tibialis anterior, electrodes were fixed lengthwise over the middle of the muscle belly. Low impedance at the skin–electrode interface was obtained (Z < 2 k{Omega}) by abrading the skin. A reference electrode was attached to the contralateral patella. The EMG signals were amplified with a bandwidth frequency ranging from 1 Hz to 1 kHz. Signals were digitized online (sampling frequency 1 kHz).

The study session started with the acquisition of 3 maximal voluntary contractions (MVC) in plantar flexion under isometric conditions. The trial with the highest value was considered to be the effective MVC. Second, isokinetic movements were realized by developing maximal concentric contractions during cycles of alternated plantar flexions and dorsiflexions at 4 angular velocities: 0.52, 1.04, 1.56, and 2.08 rad/s. The range of motion was fixed to ±0.26 rad around the reference position. For each velocity, 5 cycles were collected. Third, the elastic properties of the series elastic component were assessed by means of a quick-release technique adapted for in vivo studies (23). As in isolated muscles, the aim was to determine the stiffness of the SEC, i.e., musculotendinous stiffness. Quick-release movements were performed from the neutral position, by a sudden releasing of the moving parts of the device, while the participant maintained a submaximal plantar flexion torque equal to 25%, 50%, and 75% of MVC. Four measurements were recorded per torque level. For each trial, torque, angular displacement, and angular velocity further differentiated to obtain angular acceleration were stored for further analysis.

Finally, trials of sinusoidal oscillations were imposed on the ankle joint by using a 0.052 rad peak-to-peak angular displacement. During perturbations, the participant maintained a submaximal plantar flexion torque (20%, 40%, and 60% of MVC). Sinusoidal disturbances without participation of the participant, i.e., 0% of MVC, were also performed. Frequencies ranged from 4 Hz to 16 Hz and each trial lasted 4 s. From one trial to another, the frequency was increased by a step of 1 Hz except for frequencies higher than 12 Hz, where the steps were 2 Hz.

Data Processing
Isokinetic tests.-- A processing module allowed the researcher to select cycles of plantar flexion with a velocity plateau of correct duration and low variability in maximal torque. Parameters collected during the cycles were angular velocity, torque, and EMGs. During plantar flexion, plotting maximal concentric torque (T) against angular velocity () led to a T- relationship. As in other studies (e.g., 15), the T- curve was fitted by a logarithmic model such as:


where a and b are constants of the logarithmic function, i.e., intercept point and slope, respectively. Angular shortening velocity at low torque (T = 10% of MVC) was chosen as an index of maximal shortening velocity (VImax) in plantar flexion. Using VImax avoids the influence of the mathematical model when quantifying changes in maximal shortening velocity, since study data are often missing in the low-torque range which could led to differences in the estimation of maximal shortening velocity.

The relative decrease in maximal concentric torque, with respect to MVC, was calculated for each angular velocity (0.52, 1.04, 1.56, and 2.08 rad/s), on the basis of Equation 2:


Root mean square (RMS) values, used for expressing the power of an EMG signal, of the agonist EMGs were determined in order to verify the maximal character of the voluntary effort. RMS values of the antagonist EMGs were calculated and normalized, expressed as a percentage of the RMS value corresponding to the maximal concentric action at the same angular velocity, when the muscle acted as an agonist to get an index of coactivation.

Quick-release tests.-- Musculotendinous stiffness was measured at the beginning of the movement, when the elastic elements were supposed to recoil. Stiffness calculation was carried out within the first 20 ms of the movement (see Figure 1A). During this time lapse, no reflex changes in muscle activation (e.g., unloading reflex) were possible, which was confirmed by inspecting the recorded EMGs visually. Musculotendinous stiffness (S) was calculated as the ratio between variations in angular acceleration () and angular displacement ({Theta}) multiplied by the corresponding inertia value (I) as expressed by the formula:


In this equation, inertia is calculated and assumed to be constant. This can be verified easily by considering the transition between the static phase and the dynamic phase. At this moment, static torque (T) equals dynamic torque and acceleration is maximal (). Then:


Angular stiffness was related to the corresponding isometric torque initially exerted by the participant. The slope of the linear angular stiffness–torque relationship so obtained was defined as the stiffness index (SI) of the musculotendinous system, i.e., SIMT. Using an SI should avoid the influences of anatomical and physiological data (i.e., tendon moment arm, pennation angle, maximal voluntary contraction, cross-sectional area [CSA]) for normalizing musculotendinous stiffness, making the SI as a reliable parameter to quantify changes in musculotendinous stiffness (24).



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  		Figure 1. A, Typical raw data for quick-release movement. Changes in angular acceleration ({Delta}) and angular displacement ({Delta}{Theta}) were calculated during the first 20 ms of the movement. B, Typical data for sinusoidal perturbations. The measured parameters were torque (top) and angular displacement (bottom). The analysis only considered torque that was modulated at the driving frequency. {Phi} corresponds to the phase shift between displacement and torque

 
Sinusoidal perturbation tests.-- The measured parameters were angular displacement and torque. The subsequent analysis only considered torque that was modulated at the driving frequency in order to neglect nonlinearities (see Figure 1B) [for review, see (25)]. As in other studies [e.g., (15)], the tests were used to construct frequency–response functions (Bode diagrams) to obtain a general characterization of the musculoarticular stiffness (25). Averaged displacement-to-torque amplitude ratios (gain curve) and the phase shift between displacement and torque (phase curve) were plotted against the imposed frequencies. As shown in Figure 2, the gain diagram displays a number of characteristic features. There is an increase at low frequencies, a resonant volley at intermediate frequencies, and a decrease in gain with a slope of –40 dB/decade at high frequencies. Thus, frequency-dependent changes in gain and phase shifting reflects the classical characteristics of a mixed mechanical contribution from stiffness (K), viscosity (B), and inertia (I), of the muscle and joint system (25).



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  		Figure 2. Typical example of gain diagrams of the musculoarticular system. Data of 1 elderly woman obtained at 40% ({blacksquare}) and 60% ({square}) of maximum voluntary contraction are represented. They are adjusted by a second-order model (r2 =.985 for 40%, p <.05 and r2 =.956 for 60%, p <.05). The gain is expressed in terms of compliance

 
Using identification techniques (26), a second-order model including such parameters was adjusted to the Bode diagrams as expressed by the formula:


where {alpha} is the level of muscle activation and t is time.

Fitting by a second-order model was always satisfactory (.752 < r2 <.997, p <.05). For each level of torque, K was determined and related to torque (T). The slope of the linear K–T relationship so obtained was defined as the SI of the musculoarticular system (SIMA) in plantar flexion, and the intercept point (IP) was compared with Kp, musculoarticular stiffness in passive conditions.

Statistics
The critical level for statistical significance was set at 5%. The data are presented as means ± standard deviations (SD). Gender-related differences were analyzed with the unpaired Student's t test. In cases where the data did not meet the criteria of normality (Kolmogorov-Smirnov test, p <.05), the nonparametric Mann-Whitney rank-sum test was applied (SIMT). To compare Kp with IP, the paired Student's t test was used. Finally, a correlation coefficient (r2) was calculated for each stiffness–torque and torque–angular velocity relationship in order to determine the degree of association between variables.


   RESULTS
Top
 Abstract
 Methods
 Results
Discussion
 References
 
Contractility of Plantar Flexor Muscles
MVC of the plantar flexor muscles at 1.56 rad was found to be significantly lower for EW compared to EM (p <.05), leading to a mean difference of 17.9% (35.66 ± 4.6 vs 43.48 ± 3.9 N/m). Similarly, maximal torques during plantar flexion in isokinetic conditions were statistically lower for EW than for EM (p <.05) at each tested angular velocity, except for 0.52 rad/s. From the highest to the lowest velocity, mean maximal concentric torques ranged from 13.19 ± 1.9 to 26.03 ± 3.1 N/m for EW, and from 19.13 ± 2.2 to 30.43 ± 3.8 N/m for EM. Figure 3 illustrates typical maximal concentric torque/MVC-angular velocity relationship for 1 participant of each group. As described above, a logarithmic model was used to adjust the T-.{Theta} curves (see Eq. 1). This led to significant correlation coefficient (r2) values ranging between.887 < r2 <.993 (p <.05).



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  		Figure 3. Typical maximal torque–angular velocity relationships from isokinetic movements in plantar flexion. Data at 4 different angular velocities (0.52, 1.04, 1.56, and 2.08 rad/s) of 1 elderly woman ({square}) and 1 elderly man ({triangleup}) are represented. Moreover, maximal shortening velocity index (VImax) data are also indicated. VImax values of the woman ({blacksquare}) and the man ({blacktriangleup}), have been calculated at low torque (10% of the maximum voluntary contraction [MVC]), avoiding the influence of the mathematical model

 
As shown in Table 2, on the one hand, the relative decreases at 0.52, 1.04, and 1.56 rad/s were not significantly different between EW and EM. On the other hand, at 2.08 rad/s, the relative decreases were higher for EW than for EM (p <.05), which corresponded to a mean difference of 12.5%. Furthermore, VImax values were observed significantly lower for EW compared with EM (p <.05), leading to a mean variation of 23% (2.8 ± 0.35 vs 3.64 ± 0.25 rad/s).


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  		Table 2. Relative Maximal Torque Decreases in Plantar Flexion, During Isokinetic Tests, at Each Angular Velocity (Means ± SD).

 
Table 3 illustrates tibialis anterior index of coactivation at each angular velocity during plantar flexion contractions. No significant differences were observed between the groups.


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  		Table 3. Estimation of the Index of Coactivation of Plantar Flexion, During Isokinetic Tests (Means ± SD).

 
Musculotendinous Stiffness in Plantar Flexion
Figure 4 illustrates mean musculotendinous stiffness–torque relationships in which the best fit were linear regressions. For each individual musculotendinous stiffness–torque relationship, the correlation coefficient (r2) was found to be significant (.735 < r2 <.996, p <.05). SIMT values were significantly higher for EW compared with EM (6.73 ± 1.29 vs 4.05 ± 0.27 rad, p <.05), leading to a mean difference of 66%. In the same way, the global stiffness–torque relationships gave a global SIMT of 5.33 rad (r2 =.562, p <.05) for EW, and of 3.43 rad (r2 =.688, p <.05) for EM (see Figure 4).



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  		Figure 4. Mean musculotendinous stiffness–torque relationships from quick-release tests. Mean data and standard deviations at 3 different submaximal torques (25%, 50%, and 75% of the maximum voluntary contraction [MVC]) for the 2 groups: elderly women ({blacksquare}) and elderly men ({triangleup}) are represented. Slopes reflect musculotendinous stiffness indexes (SIMT)

 
Musculoarticular Stiffness in Plantar Flexion
In passive conditions (0% of MVC), Kp values were not significantly different between EW and EM. Mean Kp was 41.95 ± 4.46 N/m/rad for EW, and 41.74 ± 3.26 N/m/rad for EM, which was not statistically different. Furthermore, no statistical differences were found between Kp and IP for any participants (IP: 41.22 ± 4.29 for EW vs 40.28 ± 3.96 N/m/rad for EM).

Musculoarticular stiffness values, calculated at 0%, 20%, 40%, and 60% of MVC, were related to torque. The best fit were linear regressions (.834 < r2 <.999, p <.05) and the slopes were used to estimate SIMA. Figure 5 illustrates mean stiffness–torque relationships. SIMA of EW were statistically higher than those of EM (p <.05). In fact, mean SIMA was 4.73 ± 0.4 rad for EW, and 3.75 ± 0.18 rad for EM, corresponding to a difference of 26.1%. The global musculoarticular stiffness–torque relationships for the 2 populations were linear and gave a global SIMA of 4.47 rad (r2 =.842, p <.05) for the EW, of 3.7 rad (r2 =.833, p <.05) for the EM (see Figure 5).



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  		Figure 5. Mean musculoarticular stiffness–torque relationships from sinusoidal perturbations tests. Mean data and standard deviations at 4 different submaximal torques (0%, 20%, 40%, and 60% of the MVC) for the 2 groups: elderly women ({blacksquare}) and elderly men ({triangleup}) are represented. Slopes reflect musculoarticular stiffness indexes (SIMA), whereas intercepts represent passive musculoarticular stiffness (Kp)

 

   DISCUSSION
Top
 Abstract
 Methods
 Results
 Discussion
References
 
The present study was designed to quantify both human muscle and joint mechanical properties in EW and EM. Mean maximal voluntary torques in isometric and isokinetic conditions, as well as VImax were considerably lower for women than for men. Stiffness properties were quantified by 2 methods. The quick-release technique was used to determine an SI of the musculotendinous system (SIMT), which was observed to be higher for EW compared with EM. Sinusoidal perturbations were performed to quantify an SI of the musculoarticular system as a whole (SIMA), which was also greater for EW compared with EM.

Contractility of Plantar Flexor Muscles
Maximal isometric and isokinetic torques were observed to be significantly lower for EW than for EM. The large differences obtained between the 2 groups are in agreement with those found on plantar flexors of 40-to-79-year-old participants (27) and on other muscle groups such as knee extensors (2) and elbow flexors (1).

Maximal torque represents the net torque about the ankle joint (agonist minus antagonist torques). Differences in the activation level of the agonist muscles between EM and EW during maximal isometric contraction has already been evaluated by using the twitch interpolation technique (28). For instance, Vandervoort and McComas (28) have shown that most healthy EM and EW were able to maximally activate their plantar flexor muscles. During maximal concentric contraction, data on activation capacities remain scarce in the literature. Nevertheless, some studies have calculated an index of activation during elbow flexion (1,29). No differences in activation abilities have been suggested between EM and EW (1,29). According to the index of coactivation calculated during elbow flexion (1,29) and during plantar flexion in the present study, it can be hypothesized that the level of coactivation might be similar between EM and EW. Therefore, the weaker maximal isometric and isokinetic torques observed among EW might not be due to reduced activation or to higher coactivation and may be caused by other mechanisms.

One possibility can be hormonal factors, which play a role, as it has been demonstrated that strength per unit muscle mass is similar in men and premenopausal women, whereas there was a dramatic decline in strength per unit muscle mass, around the time of menopause; the latter, however, was diminished in women who used hormone replacement therapy (6).

Furthermore, the muscle strength is known to be proportional to the number of sarcomeres arranged in parallel and thus to their physiological CSA (30). In humans, muscle CSA and volume have already been investigated and estimated by ultrasonography (30) and computed tomography scanning (31). Differences in CSA have been reported between EW and EM on various muscular groups such as knee extensor muscles of 70-year-old participants (3), confirming differences in the excitable mass quantity.

In the present study, the relative decreases in maximal concentric torques with velocity calculated on the basis of Equation 2, were not statistically different between the 2 groups for velocities ranging from 0.52 rad/s to 1.56 rad/s. This leads to the assumption that muscle CSA and muscle volume are the main causes for the decline in strength. However, at 2.08 rad/s, the relative decreases were found to be higher for EW than for EM. Thus, it seems that the shape of the maximal concentric torque/MVC-velocity relationships differed for velocities higher than 2.08 rad, since significant differences in VImax were also found between EW and EM. Variations in the force–velocity relationships have been shown to be dependent on architectural parameters and fiber-type distribution (32). The internal architecture of a muscle can influence the force–velocity properties since the shortening velocity transmitted through the tendon depends on the number of sarcomeres in series with the fibers, as well as on the length of the fibers. By using ultrasonography, Kubo and colleagues (33) reported shorter gastrocnemius medialis fascicle lengths in EW compared with EM, which may partly explain the lower VImax found in EW compared with EM. Nevertheless, some studies provide evidence for differences in muscle fiber-type distribution between EW and EM. For instance, studies on the vastus lateralis indicated that EW have lower maximum unloaded shortening velocities for slow-twitch and fast-twitch fibers (34). Moreover, some authors reported that slow-twitch fibers occupied a greater area in women, whereas fast-twitch fibers occupied a greater area in men (8,9). Thus, if also present in plantar flexor muscles of EW and EM, the lower VImax of EW might be attributed to differences in architectural parameters and fiber-type distribution.

Musculotendinous Stiffness in Plantar Flexion
As in other studies (15,16), musculotendinous stiffness was expressed in angular terms, which should simplify the comparison between different groups of participants, since no assumption about the geometry of muscles and tendons is required. The present study reported for the first time gender differences in the stiffness properties of the musculotendinous system in old age, by means of a quick-release technique. Similarly to young participants (15,24), musculotendinous stiffness increased linearly with torque for each participant. SIMT values were higher for EW compared with EM.

Musculotendinous stiffness is classically separated into 2 fractions: an active fraction located in muscle fibers and a passive fraction situated notably in tendons and surrounding tissues (12). On the one hand, some results suggest that slow-twitch and fast-twitch fibers may well have different stiffness characteristics. For instance, in rats, slow-twitch fibers, such as soleus fibers, expressing the slow myosin heavy chain (MHC) isoform have been observed to be stiffer than fast-twitch fibers, such as fast extensor digitorum longus fibers, with fast MHC isoform (35). Moreover, it was demonstrated that, when a training technique increases the percentage of fast-twitch fibers in the soleus muscle of rats, its stiffness decreases (36). The opposite mechanical change (i.e., an increase in stiffness) was also associated with a relative increase in slow-twitch fibers (37). Therefore, the greater area occupied by slow-twitch fibers in women and the greater area occupied by fast-twitch fibers in men (8,9) may lead to gender differences in the elastic elements, which could influence the stiffness and partly explain the higher musculotendinous SI in EW.

Differences in musculotendinous stiffness are not only related to the active part, but the passive part of the musculotendinous system must also be considered. In humans, by using ultrasonography, the stiffness and Young's modulus in women have been found to be significantly lower than those in men, indicating that women have less stiff tendon structures compared with men (38). Consequently, the higher musculotendinous stiffness observed in EW would mainly originate from muscle fibers (i.e., active fraction). It has already been demonstrated in rat soleus muscle that a slight increase in slow-twitch fibers induces a large increase in muscle stiffness (37) after a period of endurance training.

The higher SIMT values observed in EW in the present study could cause functional differences and can be discussed in terms of improvement in muscle efficiency, i.e., force transmission via tendon directly to bone during stretch-shortening cycle (SSC) when walking. This improvement in muscle efficiency can partially compensate the lack of strength in EW compared with EM. This hypothesis can be supported by Lindle and colleagues (2) who found that the percentage of eccentric peak torque that was recovered and utilized in subsequent concentric action via SSC was greater in EW than in EM. Moreover, during the stretching phase of the SSC, the tendon which is more compliant than active muscle fibers is responsible for the greater proportion of any energy stored. During the shortening phase of the SSC, the energy stored is released. This release seems to depend mainly on the rebound resilience of the tendon and the coupling time between eccentric and concentric phases. In women, the greater rebound resilience (38) may induce a bigger energy release possibility compared with men. The coupling time might be shortened because of the higher musculotendinous stiffness in EW conducting to less energy dissipated between eccentric and concentric phases.

Musculoarticular Stiffness in Plantar Flexion
As in other studies (15,16), the mechanical behavior of the ankle joint was accurately described by a second-order model including stiffness, viscous, and inertial parameters in order to asses musculoarticular stiffness.

In passive conditions, musculoarticular stiffness values (Kp) were not significantly different between EW and EM. In fact, Kp designates a passive resistance of the ankle joint to the imposed movement when muscle is assumed to be in a resting state. Kp reflects the combined effects of passive elastic structures including skin, muscle, tendon, ligament, articular surface, and the giant protein titin. However, data in the literature concerning the interactions between the passive ankle joint structures in EM and EW are scarce.

SIMA values of EW were significantly higher than those of EM, according to the sensitivity of the sinusoidal perturbation method. In terms of musculoarticular stiffness, differences in SIMA result from structural differences in muscles, tendons, and in periarticular and intraarticular structures. SIMA values are assumed to reflect the combined effects of the musculotendinous stiffness and musculoarticular stiffness in passive conditions. On the one hand, as seen above, musculoarticular stiffness in passive conditions seems to be similar between EM and EW. On the other hand, the dependence between SIMA and active torque confirms that muscle structures contribute to stiffness measured by using sinusoidal perturbations. Thus, as shown above with the quick-release method, gender differences in musculoarticular stiffness may partly result at least from the higher musculotendinous stiffness in EW compared with EM.

In normal participants, musculoarticular stiffness is known to intervene in the postural control system by contributing to the maintenance of joint stability. On the one hand, during quiet standing (39,40), the higher musculoarticular stiffness might provide a compensatory strategy for the time delay due to sensory transduction, transmission, and processing, whereby the higher musculoarticular stiffness in EW may reduce the sensory error signal. On the other hand, according to Lakie and colleagues (41), in standing, during small perturbations, stiffness seems to act as an energy-efficient buffer. The stiffer musculoarticular system, observed in the present study, in EW would induce a less efficient buffer, i.e., stability troubles when unperceived perturbations occur. This hypothesis can be supported by Wolfson and colleagues (42) who observed that EW have impairments of balance when simultaneously deprived of visual and somatosensory inputs or during a backwards destabilization compared with EM. These researchers concluded that the higher musculoarticular stiffness in EW might explain their greater frequency of falling. In fact, falling is more common in EW than EM (43).

Conclusion
Differences between EW and EM in musculotendinous and musculoarticular stiffness during plantar flexion may be partly caused by differences in the active fraction of the musculotendinous system, i.e., differences in muscle fiber-type distribution, which is in accordance with the literature. Moreover, the present study confirms that, in old age, gender differences may exist in contractile properties under static and dynamic conditions of the ankle plantar flexor muscles, as attested by lower maximal isometric and concentric torques, and VImax for EW. Such disparities might originate from differences in contractile elements.


   Acknowledgments
 
The authors thank Yves Ballay for his technical assistance and the participants in the Better Ageing program. This work was supported by grants from the European Commission (Framework V, QLRT-2001-00323).


   Footnotes
 
Decision Editor: James R. Smith, PhD

Received December 5, 2003

Accepted February 23, 2004


   References
Top
 Abstract
 Methods
 Results
 Discussion
 References
 

  1. Valour D, Ochala J, Ballay Y, Pousson M. The influence of ageing on the force-velocity-power characteristics of human elbow flexor muscles. Exp Gerontol.. 2003;38:387-395.[Medline]
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