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Mechanomyographic and electromyographic responses of the triceps surae during maximal voluntary contractions
Journal of Electromyography and Kinesiology 13 (2003) 451–459 www.elsevier.com/locate/jelekin
Mechanomyographic and electromyographic responses of the triceps surae during maximal voluntary contractions Naokazu Miyamoto a, Shingo Oda b,∗ a
Laboratory of Human Motor Control, Graduate School of Human and Environmental Studies, Kyoto University, Sakyo-ku, Kyoto 606-8501, Japan b Laboratory of Human Motor Control, Faculty of Integrated Human Studies, Kyoto University, Kyoto 606-8501, Japan 13 November 2002
Abstract The objective of this study was to examine the effect of joint angle on the electromyogram (EMG) and mechanomyogram (MMG) during maximal voluntary contraction (MVC). Eight subjects performed maximal isometric plantar flexor torque productions at varying knee and/or ankle angles. Maximal voluntary torque, EMG, and MMG from the soleus (Sol), medial (MG) and lateral gastrocnemius (LG) muscles were measured at different joint angles. At varying knee angles, the root mean squared (rms) MMG amplitude of the MG and LG increased with knee joint extension from 60° to 180° (full extension) in steps of 30°, whereas that of the Sol was constant. At varying ankle angles, the rms-MMG of all muscles (Sol, MG, and LG) decreased with torque as ankle joint extending from 80° (10° dorsiflexion position) to 120° (30° plantar flexion position) in steps of 10°. In each case, changes in the rms-MMG of the three muscles were almost parallel to those in torque. In contrast, there were no significant differences in the rms-EMG of all muscles among all joint angles. Our data suggest that the MMG amplitudes recorded from individual muscles during MVCs can represent relative torque–angle relationships that cannot be represented by the EMG signals. 2003 Elsevier Science Ltd. All rights reserved. Keywords: Mechanomyogram; Electromyogram; Maximal voluntary contraction; Joint angle
1. Introduction Recently, mechanomyography (MMG) has been used to examine various aspects of muscle function. Investigators have suggested that the MMG reflects the motor unit (MU) activation strategy and the intrinsic mechanical activity of muscle contractions [1,28,29,45]. Many investigations [9,23,28,34,37,38,46] have utilized incremental isometric muscle contractions to determine the relationship between MMG amplitude and torque (or force). These previous investigations, however, were performed during stable contractions at separate submaximal levels at only one joint angle. Ebersole et al. [9] showed indirectly that the MMG amplitude of the rectus femoris muscle increased with flexing of the knee joint. However, they did not examine the relationship
Corresponding author. Tel.: +81-75-753-6876; fax: +81-75-7536734. E-mail address: [email protected] (S. Oda). ∗
between the maximal torque and MMG amplitude during maximal voluntary contractions (MVCs) at varying joint angles. Thus, the effect of joint angle on the MMG responses during MVCs is unclear. Based on animal studies, it is evident that the amount of isometric tension that can be produced during muscle contraction is dependent on the length at which a muscle is held [14,17]. In human experiments, torque–joint angle relationships during MVCs have been clarified [20,22,32]. The joint angle specificity for isometric torque production is due, in part, to differences in MU activation strategies [7,36,39] and/or the muscle lengthrelated overlap of actin and myosin cross-bridges [11,19,30]. Moreover, this joint angle specificity is further complicated when synergists are active that may have a nonhomogeneous input to the externally measured torque. The changes in the surface recordings of electromyographic (EMG) activity during MVCs with progressive shortening of muscle length seem to be less clear, with both increases [15,21,35] and decreases [5,31,32]. These discrepancies are possibly due to the
1050-6411/03/$ - see front matter 2003 Elsevier Science Ltd. All rights reserved. doi:10.1016/S1050-6411(03)00058-0
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many strategies of the central nervous system at varying active muscle lengths. Plantar flexor torque is mainly exerted by the triceps surae muscle group consisting of the mono-articular soleus muscle (Sol) and the bi-articular gastrocnemius muscle (medial gastrocnemius (MG) and lateral gastrocnemius (LG)). The MG and LG cross both the knee and ankle joints and therefore, by varying the knee angle, the lengths of the MG and LG are altered while the length of the Sol remains constant. As mentioned above, the EMG activity of the muscle under investigation during MVC may be influenced by the joint angle. Therefore, it is possible that the MMG amplitude–isometric torque relationships of the triceps surae muscle during MVCs may also be influenced by ankle and/or knee joint angles. Simultaneous measurements of EMG and MMG may provide insight into the relationship between the mechanical and electrical events associated with a contracting muscle. Therefore, the purpose of the present experiment was to examine the EMG and MMG responses of the triceps surae (Sol, MG, and LG) during MVCs with varying ankle and/or knee joint angles.
2. Methods 2.1. Subjects Eight male subjects (23.6 ± 1.9 years, 176.7 ± 4.5 cm, 67.3 ± 4.8 kg, mean ± SD) with no history of neuromuscular disorders volunteered for this investigation. All subjects were instructed about the experimental protocol and their informed consent was obtained. 2.2. Torque measurements Subjects lay in a prone position on a padded bed. Their right foot was tightly secured to a noncompliant footplate to ensure isometric conditions. Force was measured from beneath the ball of the foot by a load cell (LUR-A-2KNSA1; KYOWA). The axes of the ankle and footplate were aligned as closely as possible. The perpendicular distance between the load cell and axis of the footplate was used to convert force to ankle plantar flexor torque. The force transducer was zeroed to eliminate any effects of gravity and passive elasticity prior to each data collection. The present experiment was carried out in the following three conditions: constant ankle angle of 90° (Condition 1), constant knee angle of 90° (Condition 2), and constant knee angle of 180° (Condition 3). Maximal isometric contractions were performed at knee angles of 60, 90, 120, 150, and 180° (measured as the included angle between the thigh and shank) under a constant 90° ankle angle in Condition 1, at ankle angles of 80, 90, 100, 110, and 120° (measured as the internal angle
between the shank and the sole of the foot) under a constant 90° knee angle in Condition 2, and under a constant 180° knee angle in Condition 3. At each angle, the MVC was defined as the highest torque recorded in two 4 s maximal efforts. At least 1 min rest was given between MVCs within each trial at a given angle, and at least 2 min rest was given between angles to avoid fatigue. The order of testing for three conditions and the joint angles in each condition were randomized for each subject. 2.3. EMG measurements The surface EMG signals were picked up by bipolar Ag/AgCl electrodes (5 mm pick-up diameter, 30 mm inter-electrode distance), placed over the belly of the heads of the MG and LG and over the lateral aspect of the Sol along the longitudinal axis of each muscle. The reference electrode was placed over the lateral condyle of the femur for all EMG measurements. Electrode placement was preceded by abrasion of the skin surface to reduce the source impedance to less than 3 k⍀. The EMG signals were band-pass filtered (5–1000 Hz) and differentially amplified (Nihon-Kohden MEG-6108; gain: 1000 times, input impedance: ⬎100 M⍀, CMRR: ⬎80 dB). 2.4. MMG measurements For MMG sensor, an electret microphone was inserted centrally into a plastic cylinder, with an 8-mm chamber facing the skin surface. The MMG signal was detected by the specially designed sensor (QTEC, Japan) with flat frequency range of 5–1000 Hz. For each muscle (Sol, MG, and LG), one sensor was placed between the proximal and distal EMG electrodes with ring-shaped doublesided adhesive tape. The bandwidth of the MMG amplifiers was 5–100 Hz. 2.5. Signal processing The torque, EMG, and MMG signals were simultaneously and continuously stored on the hard disk of a personal computer for later analysis using a 16-bit analogue-to-digital converter (PCI-6035E, National Instruments, USA) with a sampling frequency of 1024 Hz. Recordings of the EMG, MMG, and torque signals were retrieved and analyzed for each 2 s period with a steady torque output to avoid the involvement of transient responses from rest to contraction and vice versa (Fig. 1). In addition, the digitized EMG and MMG data were processed with Hamming window function, and the fast Fourier transform was performed to obtain the mean power frequency (MPF-EMG, MMG) and the rms values of the EMG and MMG amplitude (rms-EMG, MMG), respectively.
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muscle) were used to analyze the EMG and MMG data. Follow-up analysis included one-way repeated measurements ANOVAs and Tukey post hoc comparisons. A one-way repeated measurements ANOVA and Tukey post hoc comparisons were used to analyze the torque data. In Conditions 2 and 3, separate three-way repeated measurements ANOVAs (knee angle by ankle angle by muscle) were used to analyze the EMG and MMG data. Follow-up analysis included two-way and one-way repeated measurements ANOVAs and Tukey post hoc comparison. A two-way repeated measurements ANOVA (knee angle by ankle angle) was used to analyze the torque data. Follow-up analysis included oneway repeated measurements ANOVAs and Tukey post hoc comparisons. The significance level for all comparisons was set at p ⬍ 0.05. 3. Results 3.1. Torque Fig. 2 shows the group data (mean + SD) of the maximal voluntary torque for all joint angles in each condition. In Condition 1, the plantar flexor torque increased significantly with increasing knee angle. The mean of the maximal plantar flexor torque at the 60° knee angle was about 50% of the mean maximal torque at the 180° knee angle. On the other hand, the plantar flexor torque decreased significantly with increasing ankle angle in both Conditions 2 and 3. For each ankle angle, the maximal torque in Condition 2 was significantly less than that in Condition 3. 3.2. Electromyogram
Fig. 1. Examples of raw EMG and MMG recordings of the Sol, MG, and LG and the isometric torque.
Fig. 3 shows the relationships for the rms-EMG versus joint angle for each muscle. In Condition 1, the two-way repeated measurements ANOVA revealed a significant two-way interaction. Follow-up one-way ANOVAs and Tukey post hoc comparisons showed nonsignificant differences between all knee angles for each muscle. In Conditions 2 and 3, the three-way repeated measurements ANOVA revealed a significant two-way interaction (knee angle by muscle). Follow-up one-way ANOVAs and Tukey post hoc comparisons showed nonsignificant differences between all ankle angles for each muscle and between Conditions 2 and 3 for each muscle. Fig. 4 shows the group data of the MPF-EMG of each muscle for all joint angles in each condition. In all conditions, there were no significant differences between all joint angles for each muscle.
2.6. Statistical analysis
3.3. Mechanomyogram
In Condition 1, separate two-way repeated measurements analyses of variance (ANOVAs, knee angle by
Fig. 5 shows the relationships for the rms-MMG versus joint angle for each muscle. In Condition 1, the two-
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Fig. 2. Mean (SD) isometric torque for three conditions at five different joint angles for Condition 1 (upper panel), Condition 2 (middle panel), and Condition 3 (bottom panel).
way repeated measurements ANOVA revealed a significant two-way interaction. Follow-up one-way ANOVAs and Tukey post hoc comparisons showed significant differences between knee angles for the MG (60° and 90° ⬍ 120°, 150°, and 180° knee angle and 120° ⬍ 180° knee angle) and for the LG (60° and 90° ⬍ 120°, 150°, and 180° knee angle), but no significant differences were found for the Sol. In Conditions 2 and 3, the three-way repeated measurements ANOVA
Fig. 3. Changes in root mean square amplitude of EMG (rms-EMG) obtained from the Sol (squares), MG (triangles), and LG (circles) for Condition 1, Condition 2, and Condition 3. The mean (SD) values for the eight subjects are presented. Changes in mean isometric torque (diamonds) are presented simultaneously.
resulted in significant two-way interactions (knee angle by muscle and knee angle by ankle angle). Follow-up one-way ANOVAs and Tukey post hoc comparisons showed significant differences between ankle angles for the Sol (80° ⬎ 100°, 110°, and 120° ankle angle), for the MG (80° and 90° ⬎ 100°, 110°, and 120° ankle
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Fig. 5. Changes in mean (SD) root mean square amplitude of MMG (rms-MMG) obtained from the Sol (squares), MG (triangles), and LG (circles) for Condition 1, Condition 2, and Condition 3. Changes in mean isometric torque (diamonds) are presented simultaneously.
Fig. 4. Changes in mean (SD) EMG mean power frequency (MPFEMG) of the Sol (squares), MG (triangles), and LG (circles) for Condition 1, Condition 2, and Condition 3.
angle), and the LG (80° ⬎ 100°, 110°, and 120° ankle angle) in Condition 2 and for the Sol (80° ⬎ 110° and 120° ankle angle), for the MG (80° and 90° ⬎ 120° ankle angle), and for the LG (80° ⬎ 120° ankle angle) in Condition 3. For the MG and LG in Condition 1 and the all muscles in Conditions 2 and 3, rms-MMG changed almost in parallel with torque changes at varying joint angles. Moreover, for each ankle angle, the rms-MMG of the MG and LG in Condition 2 was significantly less than that in Condition 3, respectively.
Fig. 6 shows the group data of the MPF-MMG of each muscle for all joint angles in each condition. In all conditions, there were no significant differences between all joint angles for each muscle. In all conditions, the MPFMMG of the Sol was significantly less than those of the MG and LG. In contrast, there were no significant differences between the MG and LG.
4. Discussion 4.1. Torque Maximal plantar flexor torque increased with increasing knee angle in Condition 1 and decreased with
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with muscle activation, and/or differences in the overlap of actin and myosin filaments [3,19,30,36,39,40,43]. The maximal plantar flexor torque in Condition 2 was significantly less than that in Condition 3 for each ankle angle. This was considered to be due to the bi-articular gastrocnemius muscle (MG and LG). In addition to the triceps surae, plantar flexor torque is generated by the plantaris, peroneus longus, proneus brevis, tibialis posterior, flexor hallicus longus and flexor digitorum muscles. Murray et al. [24] showed that these muscles other than the triceps surae contribute approximately 20– 30% of the total torque. Based on the data of Murray et al., Cresswell et al. [5] assumed that the torque contribution could be calculated to be at least 40% from the gastrocnemius and maximally 30% from the soleus. Furthermore, it has been shown that the torque produced by the gastrocnemius at an extreme knee flexion position is near zero [16,18]. In the present study, the maximal plantar flexor torque in Condition 2 was 50–60% of that at the same ankle angle in Condition 3. This result was close to the value estimated by Cresswell et al. [5]. 4.2. Electromyography Muscle length-dependent decreases in EMG activity have been reported in the literature for the triceps surae [5,31,32]. Cresswell et al. [5] showed that a significant decrease in the rms-EMG of the MG and LG, despite maximum voluntary effort, was observed in the shortened gastrocnemius muscle, which was possibly, in part, due to a reduced number of muscle fibers within the recording volume of the electrodes. In the present study, the rms-EMG of the MG and LG tended to increase with increasing knee angle in Condition 1, and the rms-EMG of all muscles tended to decrease with increasing ankle angle in Conditions 2 and 3, but these increases and decreases did not reach significant levels. Additionally, for the MPF-EMG, there were no significant differences between all joint angles for each muscle. Thus, based on the results of the EMG data of the present study, it seems that the activation strategies of the triceps surae would be unchanged at varying joint angles during MVCs. Fig. 6. Changes in mean (SD) MMG mean power frequency (MPFMMG) of the Sol (squares), MG (triangles), and LG (circles) for Condition 1, Condition 2, and Condition 3.
increasing ankle angle in Conditions 2 and 3. This joint angle specificity for plantar flexor torque production was consistent with previous investigations [5,32,44]. The mechanism underlying the joint angle specificity is not well known, but may be due to the selective regional activation of MUs within a particular muscle, differential activation of the plantar flexor muscles including the triceps surae, biomechanical considerations not associated
4.3. Mechanomyography Orizio [25] suggested that cross-talk between adjacent muscles may influence the recorded MMG signal. In the present study, the MMG amplitude–joint angle patterns were similar for the MG and LG. Thus, in the MMG signals, some effects of cross-talk between the MG and LG may be involved. On the other hand, the results of the present study demonstrated significant differences between the two gastrocnemius muscles and the soleus muscle in the patterns of MMG responses to varying knee joint angles. Furthermore, with the quadriceps muscle, Shinohara et al. [33] have demonstrated that appar-
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ent changes in MMG signals in the contracting muscle were few, if present at all, when extra-oscillation was added to the adjacent contracting muscle. Thus, based on these results, it seems likely that the signal detected by a sensor on a contracting muscle may reflect mainly the oscillation of the particular muscle. Relative to studies of the EMG–joint angle relationships during MVCs, fewer investigations have examined the MMG–joint angle relationship during MVCs. In the present study, the rms-MMG changed in parallel with torque and both showed the similar tendencies of change. Barry [2] reported that the MMG amplitude during maximal isometric twitches of isolated frog gastrocnemius muscles increased with stretching of muscle length up to nearly 90% of the optimal length and then decreased. Similarly, with isolated frog gastrocnemius muscles, Frangioni et al. [12] reported similar results to those shown by Barry [2]. According to a pure morphological study using ultrasonography [18], both the gastrocnemius and soleus muscles during MVC work on only the ascending limb of the force–length curve of the sarcomeres for humans [42]. Thus, based on these studies both in vitro and in vivo, it is assumed that during MVC, the maximal plantar flexor torque and the MMG amplitudes of both gastrocnemius and soleus muscles in vivo increase with increasing muscle length. This assumption is supported by the results of the present study and consistent with the results by Ebersole et al. [9]. With the rectus femoris muscle, Ebersole et al. [9] showed that the MMG amplitude increased with knee extension torque at increasing leg flexion angles from 25° to 75° below full extension. However, they did not describe the relationship between the torque and MMG amplitude during MVCs. Thus, we hypothesize that changes in the rms-MMG of each muscle at varying lengths found in the present study might be due to joint angle differences in the mechanical properties of contraction and/or slack in muscle. It has been suggested that during voluntary contractions MPF of MMG reflects the averaged firing rate of the recruited MUs [1,6,27]. Although there is still no direct evidence to prove this MMG characteristic, the suggestion has been supported by electrical stimulation studies [26,45]. In the present study, the MPF-MMG for each muscle remained unaltered during MVCs irrespective of whether its length was short or long. These results are in good agreement with the previous investigation using intramuscular electrodes [4] showing that no significant differences were found between MU firing rates at long and short muscle lengths during MVC. Based on their results, Bigland-Ritchie et al. [4] suggested that MUs firing rates during MVCs might exceed the minimum required for maximal force generation. Additionally, they showed that larger force fluctuations were found at long muscle length than at short length. During isometric contractions, there are joint angle differences
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in the mechanical properties of contraction, and contractile speed decreases and twitch duration increases at long length [8,13]. Thus, it is suggested that larger force fluctuations found at long muscle length might be due to the larger degree of synchronization between the discharges of different MUs and/or larger twitch summation of the active fibers of MUs [4]. Moreover, previous studies have indicated that MMG signals during stimulated activity reflect the force oscillation of muscle fibers [29,37,41]. Therefore, it is possible that the increases in MMG amplitude with increasing muscle length in present study may be in part due to larger mechanical oscillation on the skin overlying a contracting muscle induced by MUs synchronization and/or synchronous dimensional changes of the active fibers. However, additional research is needed to examine MUs firing rates and MUs synchronization during MVCs. An alternative hypothesis is that during isometric contractions, muscle fibers oscillations as the etiology of MMG are limited by the internal compliance of the muscle and may be reduced by stretching the muscle to more than the optimal length [2]. During MVCs, however, the human triceps surae muscles use only the ascending limb of the force–length curve. Therefore, it is possible that increasing the muscle length is not sufficient to restrict the activated muscle fibers from oscillating. In contrast, muscle fibers oscillations can occur as slack in the muscle is being reduced, so that each rms-MMG of the human triceps surae in vivo can be enhanced with increases in each muscle length. It has been reported [2,28,29,37] that high levels of muscle stiffness may restrict the muscle fiber’s ability to oscillate, thereby decreasing the MMG amplitude. Muscle stiffness is primarily a function of the number of attached cross-bridges and increases as isometric torque increases [10,11]. During MVC, however, isometric torque is determined by the number of attached crossbridges. Thus, muscle stiffness could be influenced by the number of attached cross-bridges. With regard to the results of the present study, it is possible that muscle length-related increases in muscle stiffness during MVCs were not sufficient to interfere with muscle fibers oscillations and, thereby, the MMG amplitude of each muscle increased with increasing its length. In conclusion, our results show that each MMG amplitude during MVCs changes in parallel with torque when each muscle length is altered by varying the joint angles. It seems that the MMG amplitudes recorded from individual muscles can represent the relative torque–angle relationships during MVCs that cannot be represented by the EMO signals. References [1] K. Akataki, K. Mita, M. Watakabe, K. Itoh, Mechanomyogram and force relationship during voluntary isometric ramp contrac-
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Mechanomyographic and electromyographic responses of the triceps surae during maximal voluntary contractions Naokazu Miyamoto a, Shingo Oda b,∗ a
Laboratory of Human Motor Control, Graduate School of Human and Environmental Studies, Kyoto University, Sakyo-ku, Kyoto 606-8501, Japan b Laboratory of Human Motor Control, Faculty of Integrated Human Studies, Kyoto University, Kyoto 606-8501, Japan 13 November 2002
Abstract The objective of this study was to examine the effect of joint angle on the electromyogram (EMG) and mechanomyogram (MMG) during maximal voluntary contraction (MVC). Eight subjects performed maximal isometric plantar flexor torque productions at varying knee and/or ankle angles. Maximal voluntary torque, EMG, and MMG from the soleus (Sol), medial (MG) and lateral gastrocnemius (LG) muscles were measured at different joint angles. At varying knee angles, the root mean squared (rms) MMG amplitude of the MG and LG increased with knee joint extension from 60° to 180° (full extension) in steps of 30°, whereas that of the Sol was constant. At varying ankle angles, the rms-MMG of all muscles (Sol, MG, and LG) decreased with torque as ankle joint extending from 80° (10° dorsiflexion position) to 120° (30° plantar flexion position) in steps of 10°. In each case, changes in the rms-MMG of the three muscles were almost parallel to those in torque. In contrast, there were no significant differences in the rms-EMG of all muscles among all joint angles. Our data suggest that the MMG amplitudes recorded from individual muscles during MVCs can represent relative torque–angle relationships that cannot be represented by the EMG signals. 2003 Elsevier Science Ltd. All rights reserved. Keywords: Mechanomyogram; Electromyogram; Maximal voluntary contraction; Joint angle
1. Introduction Recently, mechanomyography (MMG) has been used to examine various aspects of muscle function. Investigators have suggested that the MMG reflects the motor unit (MU) activation strategy and the intrinsic mechanical activity of muscle contractions [1,28,29,45]. Many investigations [9,23,28,34,37,38,46] have utilized incremental isometric muscle contractions to determine the relationship between MMG amplitude and torque (or force). These previous investigations, however, were performed during stable contractions at separate submaximal levels at only one joint angle. Ebersole et al. [9] showed indirectly that the MMG amplitude of the rectus femoris muscle increased with flexing of the knee joint. However, they did not examine the relationship
Corresponding author. Tel.: +81-75-753-6876; fax: +81-75-7536734. E-mail address: [email protected] (S. Oda). ∗
between the maximal torque and MMG amplitude during maximal voluntary contractions (MVCs) at varying joint angles. Thus, the effect of joint angle on the MMG responses during MVCs is unclear. Based on animal studies, it is evident that the amount of isometric tension that can be produced during muscle contraction is dependent on the length at which a muscle is held [14,17]. In human experiments, torque–joint angle relationships during MVCs have been clarified [20,22,32]. The joint angle specificity for isometric torque production is due, in part, to differences in MU activation strategies [7,36,39] and/or the muscle lengthrelated overlap of actin and myosin cross-bridges [11,19,30]. Moreover, this joint angle specificity is further complicated when synergists are active that may have a nonhomogeneous input to the externally measured torque. The changes in the surface recordings of electromyographic (EMG) activity during MVCs with progressive shortening of muscle length seem to be less clear, with both increases [15,21,35] and decreases [5,31,32]. These discrepancies are possibly due to the
1050-6411/03/$ - see front matter 2003 Elsevier Science Ltd. All rights reserved. doi:10.1016/S1050-6411(03)00058-0
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many strategies of the central nervous system at varying active muscle lengths. Plantar flexor torque is mainly exerted by the triceps surae muscle group consisting of the mono-articular soleus muscle (Sol) and the bi-articular gastrocnemius muscle (medial gastrocnemius (MG) and lateral gastrocnemius (LG)). The MG and LG cross both the knee and ankle joints and therefore, by varying the knee angle, the lengths of the MG and LG are altered while the length of the Sol remains constant. As mentioned above, the EMG activity of the muscle under investigation during MVC may be influenced by the joint angle. Therefore, it is possible that the MMG amplitude–isometric torque relationships of the triceps surae muscle during MVCs may also be influenced by ankle and/or knee joint angles. Simultaneous measurements of EMG and MMG may provide insight into the relationship between the mechanical and electrical events associated with a contracting muscle. Therefore, the purpose of the present experiment was to examine the EMG and MMG responses of the triceps surae (Sol, MG, and LG) during MVCs with varying ankle and/or knee joint angles.
2. Methods 2.1. Subjects Eight male subjects (23.6 ± 1.9 years, 176.7 ± 4.5 cm, 67.3 ± 4.8 kg, mean ± SD) with no history of neuromuscular disorders volunteered for this investigation. All subjects were instructed about the experimental protocol and their informed consent was obtained. 2.2. Torque measurements Subjects lay in a prone position on a padded bed. Their right foot was tightly secured to a noncompliant footplate to ensure isometric conditions. Force was measured from beneath the ball of the foot by a load cell (LUR-A-2KNSA1; KYOWA). The axes of the ankle and footplate were aligned as closely as possible. The perpendicular distance between the load cell and axis of the footplate was used to convert force to ankle plantar flexor torque. The force transducer was zeroed to eliminate any effects of gravity and passive elasticity prior to each data collection. The present experiment was carried out in the following three conditions: constant ankle angle of 90° (Condition 1), constant knee angle of 90° (Condition 2), and constant knee angle of 180° (Condition 3). Maximal isometric contractions were performed at knee angles of 60, 90, 120, 150, and 180° (measured as the included angle between the thigh and shank) under a constant 90° ankle angle in Condition 1, at ankle angles of 80, 90, 100, 110, and 120° (measured as the internal angle
between the shank and the sole of the foot) under a constant 90° knee angle in Condition 2, and under a constant 180° knee angle in Condition 3. At each angle, the MVC was defined as the highest torque recorded in two 4 s maximal efforts. At least 1 min rest was given between MVCs within each trial at a given angle, and at least 2 min rest was given between angles to avoid fatigue. The order of testing for three conditions and the joint angles in each condition were randomized for each subject. 2.3. EMG measurements The surface EMG signals were picked up by bipolar Ag/AgCl electrodes (5 mm pick-up diameter, 30 mm inter-electrode distance), placed over the belly of the heads of the MG and LG and over the lateral aspect of the Sol along the longitudinal axis of each muscle. The reference electrode was placed over the lateral condyle of the femur for all EMG measurements. Electrode placement was preceded by abrasion of the skin surface to reduce the source impedance to less than 3 k⍀. The EMG signals were band-pass filtered (5–1000 Hz) and differentially amplified (Nihon-Kohden MEG-6108; gain: 1000 times, input impedance: ⬎100 M⍀, CMRR: ⬎80 dB). 2.4. MMG measurements For MMG sensor, an electret microphone was inserted centrally into a plastic cylinder, with an 8-mm chamber facing the skin surface. The MMG signal was detected by the specially designed sensor (QTEC, Japan) with flat frequency range of 5–1000 Hz. For each muscle (Sol, MG, and LG), one sensor was placed between the proximal and distal EMG electrodes with ring-shaped doublesided adhesive tape. The bandwidth of the MMG amplifiers was 5–100 Hz. 2.5. Signal processing The torque, EMG, and MMG signals were simultaneously and continuously stored on the hard disk of a personal computer for later analysis using a 16-bit analogue-to-digital converter (PCI-6035E, National Instruments, USA) with a sampling frequency of 1024 Hz. Recordings of the EMG, MMG, and torque signals were retrieved and analyzed for each 2 s period with a steady torque output to avoid the involvement of transient responses from rest to contraction and vice versa (Fig. 1). In addition, the digitized EMG and MMG data were processed with Hamming window function, and the fast Fourier transform was performed to obtain the mean power frequency (MPF-EMG, MMG) and the rms values of the EMG and MMG amplitude (rms-EMG, MMG), respectively.
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muscle) were used to analyze the EMG and MMG data. Follow-up analysis included one-way repeated measurements ANOVAs and Tukey post hoc comparisons. A one-way repeated measurements ANOVA and Tukey post hoc comparisons were used to analyze the torque data. In Conditions 2 and 3, separate three-way repeated measurements ANOVAs (knee angle by ankle angle by muscle) were used to analyze the EMG and MMG data. Follow-up analysis included two-way and one-way repeated measurements ANOVAs and Tukey post hoc comparison. A two-way repeated measurements ANOVA (knee angle by ankle angle) was used to analyze the torque data. Follow-up analysis included oneway repeated measurements ANOVAs and Tukey post hoc comparisons. The significance level for all comparisons was set at p ⬍ 0.05. 3. Results 3.1. Torque Fig. 2 shows the group data (mean + SD) of the maximal voluntary torque for all joint angles in each condition. In Condition 1, the plantar flexor torque increased significantly with increasing knee angle. The mean of the maximal plantar flexor torque at the 60° knee angle was about 50% of the mean maximal torque at the 180° knee angle. On the other hand, the plantar flexor torque decreased significantly with increasing ankle angle in both Conditions 2 and 3. For each ankle angle, the maximal torque in Condition 2 was significantly less than that in Condition 3. 3.2. Electromyogram
Fig. 1. Examples of raw EMG and MMG recordings of the Sol, MG, and LG and the isometric torque.
Fig. 3 shows the relationships for the rms-EMG versus joint angle for each muscle. In Condition 1, the two-way repeated measurements ANOVA revealed a significant two-way interaction. Follow-up one-way ANOVAs and Tukey post hoc comparisons showed nonsignificant differences between all knee angles for each muscle. In Conditions 2 and 3, the three-way repeated measurements ANOVA revealed a significant two-way interaction (knee angle by muscle). Follow-up one-way ANOVAs and Tukey post hoc comparisons showed nonsignificant differences between all ankle angles for each muscle and between Conditions 2 and 3 for each muscle. Fig. 4 shows the group data of the MPF-EMG of each muscle for all joint angles in each condition. In all conditions, there were no significant differences between all joint angles for each muscle.
2.6. Statistical analysis
3.3. Mechanomyogram
In Condition 1, separate two-way repeated measurements analyses of variance (ANOVAs, knee angle by
Fig. 5 shows the relationships for the rms-MMG versus joint angle for each muscle. In Condition 1, the two-
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Fig. 2. Mean (SD) isometric torque for three conditions at five different joint angles for Condition 1 (upper panel), Condition 2 (middle panel), and Condition 3 (bottom panel).
way repeated measurements ANOVA revealed a significant two-way interaction. Follow-up one-way ANOVAs and Tukey post hoc comparisons showed significant differences between knee angles for the MG (60° and 90° ⬍ 120°, 150°, and 180° knee angle and 120° ⬍ 180° knee angle) and for the LG (60° and 90° ⬍ 120°, 150°, and 180° knee angle), but no significant differences were found for the Sol. In Conditions 2 and 3, the three-way repeated measurements ANOVA
Fig. 3. Changes in root mean square amplitude of EMG (rms-EMG) obtained from the Sol (squares), MG (triangles), and LG (circles) for Condition 1, Condition 2, and Condition 3. The mean (SD) values for the eight subjects are presented. Changes in mean isometric torque (diamonds) are presented simultaneously.
resulted in significant two-way interactions (knee angle by muscle and knee angle by ankle angle). Follow-up one-way ANOVAs and Tukey post hoc comparisons showed significant differences between ankle angles for the Sol (80° ⬎ 100°, 110°, and 120° ankle angle), for the MG (80° and 90° ⬎ 100°, 110°, and 120° ankle
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Fig. 5. Changes in mean (SD) root mean square amplitude of MMG (rms-MMG) obtained from the Sol (squares), MG (triangles), and LG (circles) for Condition 1, Condition 2, and Condition 3. Changes in mean isometric torque (diamonds) are presented simultaneously.
Fig. 4. Changes in mean (SD) EMG mean power frequency (MPFEMG) of the Sol (squares), MG (triangles), and LG (circles) for Condition 1, Condition 2, and Condition 3.
angle), and the LG (80° ⬎ 100°, 110°, and 120° ankle angle) in Condition 2 and for the Sol (80° ⬎ 110° and 120° ankle angle), for the MG (80° and 90° ⬎ 120° ankle angle), and for the LG (80° ⬎ 120° ankle angle) in Condition 3. For the MG and LG in Condition 1 and the all muscles in Conditions 2 and 3, rms-MMG changed almost in parallel with torque changes at varying joint angles. Moreover, for each ankle angle, the rms-MMG of the MG and LG in Condition 2 was significantly less than that in Condition 3, respectively.
Fig. 6 shows the group data of the MPF-MMG of each muscle for all joint angles in each condition. In all conditions, there were no significant differences between all joint angles for each muscle. In all conditions, the MPFMMG of the Sol was significantly less than those of the MG and LG. In contrast, there were no significant differences between the MG and LG.
4. Discussion 4.1. Torque Maximal plantar flexor torque increased with increasing knee angle in Condition 1 and decreased with
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with muscle activation, and/or differences in the overlap of actin and myosin filaments [3,19,30,36,39,40,43]. The maximal plantar flexor torque in Condition 2 was significantly less than that in Condition 3 for each ankle angle. This was considered to be due to the bi-articular gastrocnemius muscle (MG and LG). In addition to the triceps surae, plantar flexor torque is generated by the plantaris, peroneus longus, proneus brevis, tibialis posterior, flexor hallicus longus and flexor digitorum muscles. Murray et al. [24] showed that these muscles other than the triceps surae contribute approximately 20– 30% of the total torque. Based on the data of Murray et al., Cresswell et al. [5] assumed that the torque contribution could be calculated to be at least 40% from the gastrocnemius and maximally 30% from the soleus. Furthermore, it has been shown that the torque produced by the gastrocnemius at an extreme knee flexion position is near zero [16,18]. In the present study, the maximal plantar flexor torque in Condition 2 was 50–60% of that at the same ankle angle in Condition 3. This result was close to the value estimated by Cresswell et al. [5]. 4.2. Electromyography Muscle length-dependent decreases in EMG activity have been reported in the literature for the triceps surae [5,31,32]. Cresswell et al. [5] showed that a significant decrease in the rms-EMG of the MG and LG, despite maximum voluntary effort, was observed in the shortened gastrocnemius muscle, which was possibly, in part, due to a reduced number of muscle fibers within the recording volume of the electrodes. In the present study, the rms-EMG of the MG and LG tended to increase with increasing knee angle in Condition 1, and the rms-EMG of all muscles tended to decrease with increasing ankle angle in Conditions 2 and 3, but these increases and decreases did not reach significant levels. Additionally, for the MPF-EMG, there were no significant differences between all joint angles for each muscle. Thus, based on the results of the EMG data of the present study, it seems that the activation strategies of the triceps surae would be unchanged at varying joint angles during MVCs. Fig. 6. Changes in mean (SD) MMG mean power frequency (MPFMMG) of the Sol (squares), MG (triangles), and LG (circles) for Condition 1, Condition 2, and Condition 3.
increasing ankle angle in Conditions 2 and 3. This joint angle specificity for plantar flexor torque production was consistent with previous investigations [5,32,44]. The mechanism underlying the joint angle specificity is not well known, but may be due to the selective regional activation of MUs within a particular muscle, differential activation of the plantar flexor muscles including the triceps surae, biomechanical considerations not associated
4.3. Mechanomyography Orizio [25] suggested that cross-talk between adjacent muscles may influence the recorded MMG signal. In the present study, the MMG amplitude–joint angle patterns were similar for the MG and LG. Thus, in the MMG signals, some effects of cross-talk between the MG and LG may be involved. On the other hand, the results of the present study demonstrated significant differences between the two gastrocnemius muscles and the soleus muscle in the patterns of MMG responses to varying knee joint angles. Furthermore, with the quadriceps muscle, Shinohara et al. [33] have demonstrated that appar-
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ent changes in MMG signals in the contracting muscle were few, if present at all, when extra-oscillation was added to the adjacent contracting muscle. Thus, based on these results, it seems likely that the signal detected by a sensor on a contracting muscle may reflect mainly the oscillation of the particular muscle. Relative to studies of the EMG–joint angle relationships during MVCs, fewer investigations have examined the MMG–joint angle relationship during MVCs. In the present study, the rms-MMG changed in parallel with torque and both showed the similar tendencies of change. Barry [2] reported that the MMG amplitude during maximal isometric twitches of isolated frog gastrocnemius muscles increased with stretching of muscle length up to nearly 90% of the optimal length and then decreased. Similarly, with isolated frog gastrocnemius muscles, Frangioni et al. [12] reported similar results to those shown by Barry [2]. According to a pure morphological study using ultrasonography [18], both the gastrocnemius and soleus muscles during MVC work on only the ascending limb of the force–length curve of the sarcomeres for humans [42]. Thus, based on these studies both in vitro and in vivo, it is assumed that during MVC, the maximal plantar flexor torque and the MMG amplitudes of both gastrocnemius and soleus muscles in vivo increase with increasing muscle length. This assumption is supported by the results of the present study and consistent with the results by Ebersole et al. [9]. With the rectus femoris muscle, Ebersole et al. [9] showed that the MMG amplitude increased with knee extension torque at increasing leg flexion angles from 25° to 75° below full extension. However, they did not describe the relationship between the torque and MMG amplitude during MVCs. Thus, we hypothesize that changes in the rms-MMG of each muscle at varying lengths found in the present study might be due to joint angle differences in the mechanical properties of contraction and/or slack in muscle. It has been suggested that during voluntary contractions MPF of MMG reflects the averaged firing rate of the recruited MUs [1,6,27]. Although there is still no direct evidence to prove this MMG characteristic, the suggestion has been supported by electrical stimulation studies [26,45]. In the present study, the MPF-MMG for each muscle remained unaltered during MVCs irrespective of whether its length was short or long. These results are in good agreement with the previous investigation using intramuscular electrodes [4] showing that no significant differences were found between MU firing rates at long and short muscle lengths during MVC. Based on their results, Bigland-Ritchie et al. [4] suggested that MUs firing rates during MVCs might exceed the minimum required for maximal force generation. Additionally, they showed that larger force fluctuations were found at long muscle length than at short length. During isometric contractions, there are joint angle differences
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in the mechanical properties of contraction, and contractile speed decreases and twitch duration increases at long length [8,13]. Thus, it is suggested that larger force fluctuations found at long muscle length might be due to the larger degree of synchronization between the discharges of different MUs and/or larger twitch summation of the active fibers of MUs [4]. Moreover, previous studies have indicated that MMG signals during stimulated activity reflect the force oscillation of muscle fibers [29,37,41]. Therefore, it is possible that the increases in MMG amplitude with increasing muscle length in present study may be in part due to larger mechanical oscillation on the skin overlying a contracting muscle induced by MUs synchronization and/or synchronous dimensional changes of the active fibers. However, additional research is needed to examine MUs firing rates and MUs synchronization during MVCs. An alternative hypothesis is that during isometric contractions, muscle fibers oscillations as the etiology of MMG are limited by the internal compliance of the muscle and may be reduced by stretching the muscle to more than the optimal length [2]. During MVCs, however, the human triceps surae muscles use only the ascending limb of the force–length curve. Therefore, it is possible that increasing the muscle length is not sufficient to restrict the activated muscle fibers from oscillating. In contrast, muscle fibers oscillations can occur as slack in the muscle is being reduced, so that each rms-MMG of the human triceps surae in vivo can be enhanced with increases in each muscle length. It has been reported [2,28,29,37] that high levels of muscle stiffness may restrict the muscle fiber’s ability to oscillate, thereby decreasing the MMG amplitude. Muscle stiffness is primarily a function of the number of attached cross-bridges and increases as isometric torque increases [10,11]. During MVC, however, isometric torque is determined by the number of attached crossbridges. Thus, muscle stiffness could be influenced by the number of attached cross-bridges. With regard to the results of the present study, it is possible that muscle length-related increases in muscle stiffness during MVCs were not sufficient to interfere with muscle fibers oscillations and, thereby, the MMG amplitude of each muscle increased with increasing its length. In conclusion, our results show that each MMG amplitude during MVCs changes in parallel with torque when each muscle length is altered by varying the joint angles. It seems that the MMG amplitudes recorded from individual muscles can represent the relative torque–angle relationships during MVCs that cannot be represented by the EMO signals. References [1] K. Akataki, K. Mita, M. Watakabe, K. Itoh, Mechanomyogram and force relationship during voluntary isometric ramp contrac-
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