Modulation of Stretch Reflexes During Imposed Walking Movements of the Human Ankle

Robert E. Kearney,1 Mireille Lortie,1 and Richard B. Stein2

 1Department of Biomedical Engineering, McGill University, Montreal, Quebec H3A 2B4; and  2Division of Neuroscience, University of Alberta, Edmonton, Alberta T6G 2S2, Canada


    ABSTRACT
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES

Kearney, Robert E., Mireille Lortie, and Richard B. Stein. Modulation of stretch reflexes during imposed walking movements of the human ankle. Our overall objectives were to examine the role of peripheral afferents from the ankle in modulating stretch reflexes during imposed walking movements and to assess the mechanical consequences of this reflex activity. Specifically we sought to define the changes in the electromyographic (EMG) and mechanical responses to a stretch as a function of the phase of the step cycle. We recorded the ankle position of a normal subject walking on a treadmill at 3 km/h and used a hydraulic actuator to impose the same movements on supine subjects generating a constant level of ankle torque. Small pulse displacements, superimposed on the simulated walking movement, evoked stretch reflexes at different phases of the cycle. Three major findings resulted: 1) soleus reflex EMG responses were influenced strongly by imposed walking movements. The response amplitude was substantially smaller than that observed during steady-state conditions and was modulated throughout the step cycle. This modulation was qualitatively similar to that observed during active walking. Because central factors were held constant during the imposed walking experiments, we conclude that peripheral mechanisms were capable of both reducing the amplitude of the reflex EMG and producing its modulation throughout the movement. 2) Pulse disturbances applied from early to midstance of the imposed walking cycle generated large reflex torques, suggesting that the stretch reflex could help to resist unexpected perturbations during this phase of walking. In contrast, pulses applied during late stance and swing phase generated little reflex torque. 3) Reflex EMG and reflex torque were modulated differently throughout the imposed walking cycle. In fact, at the time when the reflex EMG response was largest, the corresponding reflex torque was negligible. Thus movement not only changes the reflex EMG but greatly modifies the mechanical output that results.


    INTRODUCTION
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES

The stretch reflex has been the topic of intense study for many years (Davidoff 1992; Liddell and Sherrington 1924; Matthews 1970). It is mediated at least in part by monosynaptic connections from primary muscle spindle afferents to alpha -motoneurons. Despite the apparent simplicity of the underlying pathways, the role of the stretch reflex in the control of posture and movement remains unclear. One reason for this is the difficulty in separating the mechanical consequences of the stretch reflex from those due to intrinsic mechanics. In addition, there is growing evidence that the gain of the stretch reflex is generally lower during movement than at rest and is modulated systematically throughout cyclic activities (reviewed by Brooke et al. 1997). Consequently findings about stretch reflexes during postural conditions, the most commonly studied situation, may have limited applicability to reflex function during movement.

Most evidence regarding the modulation of reflexes has come from studies of the H reflex, which is evoked by the electrical stimulation of group Ia afferents. H reflexes are modulated strongly during rhythmic activities such as stepping (Brooke et al. 1995a; Crenna and Frigo 1987), walking (Brooke et al. 1991; Capaday and Stein 1986), running (Capaday and Stein 1987), hopping, and pedaling (Brooke et al. 1993; Collins et al. 1993; McIlroy et al. 1992). There are two components to this modulation: a tonic inhibition, correlated with the velocity of limb movement (Collins et al. 1993; McIlroy et al. 1992), and a phasic modulation in which the reflex response changes systematically throughout the cycle. For example, during walking, reflex excitability is larger during the stance phase than during the swing phase (Capaday and Stein 1986; Crenna and Frigo 1987; Yang and Stein 1990). Brooke and colleagues have presented a variety of evidence indicating that this reflex inhibition arises from peripheral input, probably from muscle afferents acting presynaptically (Brooke et al. 1997). However, there is also evidence pointing to a role for central factors (Yang and Whelan 1993).

H reflexes provide a convenient means of assessing the excitability of the neural pathways underlying the stretch reflex. However, the functional significance of changes in H reflexes is difficult to interpret because other factors could modify the effectiveness of the stretch reflex. First, fusimotor drive could modulate the sensitivity of muscle spindle receptors independently of motor neuron excitability. This possibility has been explored by comparing electromyographic (EMG) responses evoked by sudden stretch and by electrical stimulation during walking. Results to date have not been conclusive; one study reported significant differences in behavior between H and stretch reflexes (Sinkjaer et al. 1996) while another (Yang et al. 1991) concluded that the two responses behaved in a qualitatively similar manner. Second, the mechanical consequences of this reflex EMG activity are likely to vary strongly throughout the cycle due to the nonlinear dependence of muscle on length and velocity. Unfortunately, for technical reasons, reflex torque has not been measured during normal locomotion and so the importance of these effects is not known.

The objectives of this study were to examine the role of peripheral afferents from the ankle in the modulation of stretch reflexes during imposed walking movements and to assess the mechanical consequences of this reflex activity. In particular, we sought to define the changes in the EMG and mechanical responses to a standard stretch input as a function of the phase of the step cycle. To do this, we recorded the ankle position of a normal subject walking on a treadmill at 3 km/h and then used a hydraulic actuator to impose these walking movements while the subject remained supine and generated a constant level of ankle torque. Thus central factors were held as constant as possible so that any reflex modulation that occurred could be attributed to peripheral effects. Stretch reflex behavior was tested at 10 different points in the cycle by applying small, rapid pulse displacements. The amplitude of the reflex EMG was reduced during movement in comparison with responses evoked at matched static conditions. Moreover the amplitude of the reflex EMG responses was modulated throughout the imposed walking cycle in a manner similar to that seen in normal walking, suggesting that afferents from the ankle play an important role in modulating the reflex. Substantial reflex torques were generated during the first half of the stance phase. Indeed their amplitude was comparable with the responses observed under matched static conditions. However, pulses applied during the later portion of stance generated no significant torque, although the EMG responses were large. This dissociation between the behavior of the reflex EMG and reflex torque demonstrates that muscle dynamics play a key role in determining the functional significance of stretch reflex activity.

The results of this work have been communicated at conferences (Kearney et al. 1996; Lortie et al. 1997).


    METHODS
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ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES

Subjects

Six subjects (5 male, 1 female) between the ages of 20 and 39 yr with no history of neuromuscular disease were studied. All subjects gave informed consent to the experimental procedures, which had been reviewed and approved by McGill University's Research Ethics Board.

Data acquisition

Subjects lay supine on the experimental table with their foot fixed to the pedal of an electro-hydraulic actuator by a rigid custom-fitted boot (Morier et al. 1990). The knee joint was held in a slightly flexed position (~20°) by sandbags and a knee strap.

Ankle position was measured with a plastic-film potentiometer mounted in the actuator. Torque was recorded using a torque transducer mounted in series with the subject's ankle; the stiffness of the transducer (50 kNm/rad) was much greater than that of the ankle. Surface EMGs were recorded from tibialis anterior (TA) and soleus muscle using bipolar surface electrodes. The TA electrodes were placed ~2.5 cm apart over the belly of the muscle, approximately one-third of the distance from the knee to the ankle. The soleus electrodes were placed 2.5 cm apart on the soleus muscle slightly lateral to the midline just distal to the heads of the gastrocnemius. With custom-built processing electronics (Perreault et al. 1993), EMG signals were first amplified differentially (gain of 1,000 or 10,000), high-pass filtered with a first-order filter (1-Hz cutoff), and then full-wave rectified. All signals were anti-alias filtered at 200 Hz using eight-pole Bessel filters and sampled at 1 kHz by a 16-bit A/D converter.

Experimental procedures

ZERO POSITION. The range of motion of the ankle was determined by passively moving the foot with the actuator power off. Safety stops were adjusted to prevent the actuator range from exceeding the subject's range of movement. The ankle then was moved to a position 0.3 rad from full dorsiflexion, and baseline measurements were made. Subsequently all measurements of ankle angle were made with reference to this position with dorsiflexing displacements being taken as positive and plantarflexing displacements as negative.

MAXIMUM VOLUNTARY CONTRACTION. With the ankle in the zero position, subjects were instructed to make a maximum voluntary contraction (MVC) in the plantarflexing direction for 5 s. Torque and EMGs were recorded throughout the contraction and for 5 s before and after the contraction. EMG records were inspected to ensure that EMG gains were adequate and that cross talk was not excessive. The torque record was inspected to ensure that a plateau was achieved, consistent with the subject having generated an MVC (Kroemer and Marras 1980). If so, the maximum torque recorded was used as a measure of the MVC; if not, the MVC trial was repeated.

IMPOSED WALKING MOVEMENTS. Imposed walking movements were generated by using a record of the average ankle position during walking as the command input to the ankle actuator. This input signal was obtained in an initial set of experiments in which ankle position was recorded as a subject walked on a treadmill at 3 km/h. Sixty cycles of position data were recorded, aligned to a common reference point corresponding to the beginning of the stance phase (heel strike) and ensemble averaged. The actuator's dynamic response and torque capabilities were such that each cycle of the imposed ankle movement was almost identical to the command input (the variance accounted for between the desired and actual ankle position was always >99.8%).

Subjects were instructed to maintain a constant, voluntary, plantarflexing contraction at a level equal to 5% of their MVC aided by a visual display of the target and feedback torques. At this level of contraction, most of the EMG and torque were produced by the soleus muscles. The feedback torque was obtained by low-pass filtering ankle torque at 0.2 Hz using an eight-pole Bessel filter. This low cutoff frequency was chosen to eliminate intracycle variations in torque arising from intrinsic mechanisms. Subjects rapidly learned to keep the mean torque nearly constant from cycle to cycle.

Torque rather than EMG was used as the feedback signal because the EMG responses to pulse displacements were much larger, relative to the background activity, than the torque responses. As a result, the EMG feedback signal fluctuated much more than the torque feedback, making it more difficult for the subject to maintain a constant level of activity.

Every third or fourth cycle, a small pulse displacement, lasting ~40 ms, was superimposed on the imposed walking movements to elicit a stretch reflex (see Fig. 1). Pulses were applied, in random order, at 10 locations equally spaced throughout the walking cycle, as shown in Fig. 3. Pulse amplitude (0.0375 rad for 2 subjects and 0.05 for 4 subjects) was selected in initial trials, under static conditions, by determining the pulse amplitude which maximized the reflex torque and did not break a significant number of cross bridges in the background contraction. Breaking a large number of cross bridges is undesirable because it results in a drop in torque immediately after the pulse, making it difficult to determine the true magnitude of the reflex torque.



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Fig. 1. Ensemble averages of position (A), rectified soleus electromyogram (EMG; B), and torque (C) for a typical pulse location. Two cycles are shown with the time corresponding to heelstrike (- - -). Zero value for position was defined to be 0.3 rad from full dorsiflexion. Dorsiflexing positions and torques are taken as positive and the soleus EMG has been rectified in the positive direction. Pulse superimposed on the position input during the 2nd cycle evoked large EMG and torque responses.

Sampling began one cycle before the application of the pulse so that each data record comprised a control cycle followed by one with a perturbation, as illustrated in Fig. 1, top. The mean value of the torque in the control cycle was computed and the trial discarded if it deviated from the desired value by >1 Nm. Acquisition continued until 10 trials were obtained for each position.

STATIC TRIALS. The static position dependence of the stretch reflex was examined by recording the responses to 10 pulse displacements by themselves, with no superimposed walking, at different positions through the range of motion. The mean voluntary torque level and pulse amplitudes were matched to those used during the imposed walking trials.

STATIONARITY. These experiments were quite time consuming; acquisition of walking data could take > 20 min, depending on how well the subject was able to maintain the desired torque level. Performing the static trials at the different positions required a similar time. Fatigue was not usually a problem due to the low level of the tonic contraction. Nevertheless, we were concerned that there might be time-dependent changes in the stretch reflex. To test for this, control trials, in which pulses were applied at the zero position under static conditions, were done at regular intervals throughout the experiment. Small (20-30%), but statistically significant, changes in the responses were observed occasionally. To prevent this from biasing our results, we randomized the order in which trials were done. For static trials, the various positions were tested in random order; for the walking trials, the location of each pulse was selected randomly. Furthermore the static trials were done prior to the walking trials for three subjects and after the walking trials for the other three.

Analysis

AVERAGING. The 10 trials acquired at each position in the static paradigm were ensemble averaged prior to analysis. The 10 perturbed cycles recorded at each stimulus location during imposed walking movements were ensemble averaged to form an average perturbed trial. The average control cycle was computed by aligning the control cycles from all trials (10 for each of the 10 positions) to a common reference point and then ensemble averaging. This procedure, which increased the number of responses in the ensemble to 100, resulted in a noticeably better signal than simply using the 10 control trials acquired at each stimulus location. The incremental EMGs and torques, evoked by the pulse at each location, were determined by taking the difference between the average perturbed and control cycles, after first aligning both to a common reference point.

PULSE RESPONSE. Pulse responses, for both static and walking paradigms, were analyzed with methods similar to those used by Stein and Kearney (1995). Briefly, the reflex EMG amplitude was defined as the maximum absolute value of the incremental EMG signal from 30 to 70 ms after the onset of the pulse. Similarly the reflex torque amplitude was defined as the maximum absolute value of the incremental torque from 100 to 300 ms after pulse onset.


    RESULTS
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES

Pulse responses

Figure 1 shows ensemble average data from one subject for a typical pulse location. It comprises two complete cycles of imposed walking movements; the time heel strike occurred in the treadmill experiment is marked (- - -). The ankle position (Fig. 1A) started with a brief plantarflexion (downward deflection), from the time of heel strike until the foot was flat on the ground. Then there is a period of steady dorsiflexion (upward deflection) lasting ~500 ms, corresponding to much of the stance phase. The ankle then plantarflexed rapidly, as first the heel and then the toe left the ground, and returned to the original plantarflexed position in ~200 ms. The ankle then dorsiflexed slightly and remained at a nearly constant position for the remaining 250 ms of the swing phase. EMG activity in the soleus muscle (Fig. 1B) was generally low although sporadic activity was present throughout the cycle for this subject. Ankle torque (Fig. 1C) had a peak-to-peak amplitude of ~10 Nm and varied reciprocally to position as would be expected from joint visco-elasticity. The pulse perturbation in this trial can be seen in the second cycle starting ~235 ms after heel strike. Although the pulse amplitude (0.05 rad) was small relative to the peak-to-peak amplitude of the walking movement (0.45 rad), large responses are apparent in both the soleus EMG and the torque records.

The responses evoked by the pulse are seen more clearly in Fig. 2, which shows the incremental signals computed by taking the difference between the averaged control and perturbed trials. The pulse input (Fig. 2A) was a rapid dorsiflexing movement with amplitude of 0.05 rad, a maximum velocity of 3.7 rad/s, and a duration of 42 ms. The reflex soleus EMG response (Fig. 2B) was dominated by a single, phasic burst of activity having a latency of 45 ms. The TA EMG record is not shown because the only activity observed was small and likely due to cross talk. The torque response (Fig. 2C) comprised two distinct components: a short-latency component associated with intrinsic mechanisms (the inertial, viscous and elastic properties of joint and muscle) and a longer latency transient due to reflex activation. The two components are separated clearly in this record because the pulse displacement was completed just before the reflex torque began. The responses closely resembled those described previously for pulses applied under static conditions or with superimposed random perturbations (Stein and Kearney 1995).



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Fig. 2. Incremental responses obtained by subtracting the average unperturbed cycle from the perturbed cycle shown in Fig. 1. A: position. B: soleus EMG. C: torque.

Modulation through the walking cycle

The reflex EMG responses evoked by pulses applied at 10 different locations throughout the simulated walking cycle are shown in Fig. 3A. The top curve shows the 10 pulses superimposed on the basic walking pattern. The 10 reflex EMG responses are shown below. For clarity, each trace was offset vertically and only 350 ms of data are shown, starting 100 ms before the onset of the pulse. The reflex EMG amplitude was low at heel contact, increased progressively throughout the simulated stance phase, and reached its maximum value near full dorsiflexion. It dropped nearly to zero as soon as the ankle began to plantarflex and remained low throughout the simulated swing phase.



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Fig. 3. Modulation of reflex response throughout the imposed walking cycle. A, top: average ankle position with the 10 pulse displacements superimposed. Bottom: incremental EMG responses to all 10 perturbations for 1 subject. For clarity, only a 350-ms segment of each response is shown and the traces are offset vertically. Heel strike occurs at time 0. B: incremental torque responses in the same format as A. Inset: initial portion of the 10 torque responses superimposed on an expanded time scale.

Figure 3B shows the variation of the torque response throughout the walking cycle. As with the EMG response, the reflex torques were low at heel strike and increased as the ankle dorsiflexed during the early and middle parts of the stance phase. However, pulses applied late in the stance phase evoked little if any reflex torque despite generating large EMG responses. Pulses applied during the swing phase generated no significant torque response consistent with the absence of an EMG response.

Figure 3B, inset, shows an enlarged view of the early part of the 10 torque responses superimposed and demonstrates two important points. First, the early responses (0-25 ms), due primarily to inertial effects, are nearly identical. This indicates that the mechanical coupling between the foot and the actuator was much the same for all 10 responses. Any change in coupling would be expected to have dramatic effects on the early response. Second, the torque returns to baseline at the end of the pulse displacement (~75 ms), demonstrating that the muscle retains its ability to generate force. If this was not the case, due to the breakage of cross bridges for example, the torque would return to a value above the baseline. (Note that because these are incremental responses the 0 level corresponds to the constant plantarflexing contraction).

Group responses

There was considerable intersubject variation in the magnitude of the reflex responses and MVC. This variability is summarized in Table 1, which lists the MVC for each subject as well as the maximum reflex torque recorded in the "imposed-walking" and "static-control" paradigms. Maximum reflex torques, as a percentage of MVC, were comparable in size with those we described earlier (Stein and Kearney 1995). The reflex torque magnitude was not correlated with that of the MVC, so normalizing the reflex torque to MVC did not reduce the subject-to-subject variability.


                              
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Table 1. Reflex and MVC torques

In contrast, the modulation of the reflex EMG and torque throughout the imposed walking cycle was very consistent for all subjects. This is illustrated in Fig. 4, which shows the data from all six subjects plotted separately. In all subjects, the reflex EMG increased steadily to reach a peak toward the end of the simulated stance phase. The reflex torque also increased to a peak, but this peak occurred near the middle of the simulated stance phase. This clear discrepancy between the timing of the peaks in EMG and torque has not been reported before to our knowledge.



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Fig. 4. Reflex EMG (A) and torque (C) plotted as functions of time for all 6 subjects in this study. Right: reflex EMG (B) and torque (D) observed under static conditions at matched torque levels and positions are plotted as functions of the time when the corresponding position occurred in the step cycle.

Static position

Previous results have shown that the stretch reflex changes dramatically with static joint position; the response to a standard pulse displacement increases greatly as the ankle is dorsiflexed progressively (Stein and Kearney 1995). Consequently the reflex modulation that was observed during these imposed walking movements might simply reflect the static position dependence. To assess this possibility, we compared the pattern of reflex modulation observed during imposed walking movement with that observed at equivalent positions under static conditions. For this purpose, pulse displacements, identical to those used in the walking experiments, were applied under postural conditions at positions spanning the range of motion. Subjects were required to generate the same mean torque during these static trials as they did during the imposed walking. However, the positions at which pulses were applied during the static trials were not precisely the same as those where pulses were applied during imposed walking. Consequently we used linear interpolation to estimate the responses, under static conditions, at each of the 10 positions tested during the imposed walking experiments. These values are plotted in Fig. 4 as a function of time to facilitate comparison with results obtained during imposed walking.

Three points are evident: 1) the reflex EMG was consistently much larger under static conditions at the same muscle lengths; 2) the region of high reflex EMG and torque is much broader under static conditions than under dynamic conditions; and 3) there is no obvious discrepancy in the positions of the peak EMG and torque under static conditions.

Group means

Because these new findings were consistent in all subjects studied, group means of the amplitudes of the reflex EMG and torque were computed to summarize the overall pattern of reflex modulation. The amplitudes of the EMG and torque responses for each of the six subjects were normalized to their maximum values during imposed walking, and then group means and standard errors were computed at each point in the cycle. As was evident from the individual values, the group mean reflex EMG, indicated in Fig. 5B (), was low at heel strike, increased progressively throughout the simulated stance phase, rapidly decreased as the ankle began to plantarflex, and remained low throughout the simulated swing phase. The standard errors were small, reflecting our observation that the data from all subjects followed the same general pattern.



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Fig. 5. Group average responses in imposed walking and static trials. A: ankle position with the 10 pulse displacements superimposed. B: mean ± 1 SE of the normalized reflex EMG for simulated walking () and related static trials (). C: mean ± 1 SE normalized reflex torque for simulated walking () and related static trials ().

The group mean reflex torque, shown as circles in Fig. 5C, increased early in stance and reached a maximum corresponding to midstance. It then decreased as the simulated stance phase was completed, although the EMG response continued to increase. Indeed, the pulse applied at the peak of dorsiflexion (600 ms) evoked the largest EMG response but generated almost no reflex torque. As was shown in Fig. 4 for individual subjects, little or no reflex torque was evoked during the portion of the imposed movement corresponding to toe-off and swing phase.

Figure 5B also shows the group mean EMG responses for the static trials () obtained by normalizing each subject's data to the maximum value observed during imposed walking, after interpolating and averaging. As shown for all subjects individually, the reflex EMG was substantially smaller during the imposed walking movement than under static conditions. Indeed, sequential t-tests confirmed that the static responses were significantly larger (P < 0.05) at all points except that at 800 ms, where neither value was significantly greater than zero. In addition, the shapes of the two curves were quite different, indicating that the modulation of the reflex EMG during walking does not depend simply on static position.

Figure 5C shows the same comparison for reflex torques. During the period corresponding to the first half of stance, the average torque responses for the imposed walking and static trials were not significantly different (P < 0.05). Reflex torques were significantly greater under static conditions than during the imposed walking at all other points in the cycle, except at 800 ms where the reflex torque was near zero for both conditions.

Velocity effects

The pulse perturbation might have varied systematically throughout the cycle due to interactions between the walking and pulse stimuli or because of actuator limitations. Although variations in the amplitude of the pulse perturbation were small, we felt that in view of the velocity-dependent characteristics of the stretch reflex we should examine the maximum pulse velocity (the difference between the control and perturbed trials) at the 10 positions. Figure 6A () shows the group mean pulse velocity as a function of latency. There is a small but significant decrease at 600 ms, where the EMG response was maximal, but no differences elsewhere. Consequently the incremental velocity cannot account for the modulation we observed.



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Fig. 6. A: velocity of the imposed walking movement (down-triangle), velocity of the pulse perturbations (), and the net velocity (open circle ). B: background EMG generated by each of the 6 subjects over the simulated walking cycle. Mean values were measured for a 100-ms period centered on the times indicated.

The velocity of the imposed walking movements varied from -2 to +1 rad/s as shown in Fig. 6A (triangle ------triangle ). Pulses were superimposed on these movements so that the net velocity, the sum of the walking and pulse velocities (open circle ), was more variable than the pulse velocity. It was particularly low at 700 ms because at this phase the velocity of walking was opposite to that of the pulse. The small reflex response observed at this point may be due to the low net velocity. However, elsewhere in the cycle there was little correlation between the reflex responses and net velocity. Net velocity remained nearly constant throughout most of the simulated stance phase (100-500 ms), whereas the reflex response increased progressively. Indeed the reflex response reached a maximum at 600 ms despite a drop in the absolute velocity. Clearly the variation in the net velocity cannot account for the large modulation of the reflex EMG and torque.

Voluntary activation

The reflex modulation also might arise from systematic variations in the central drive. To test this possibility, we averaged the background EMG activity in the control trials for the 100-ms intervals centered about the times of pulse application. Although subjects did their best to maintain a constant torque level, there were some systematic fluctuations in the EMG of individual subjects. For example, one subject (diamond ) in Fig. 6B had two peaks of EMG, one in midstance and one in swing. Another subject (open circle ) had one peak in late stance. Others showed no obvious peaks, yet all showed the same pattern of reflex EMG in Fig. 4. Thus although some subjects were not able to maintain the level of voluntary EMG precisely constant, the random variations that occurred were not correlated with the systematic variation in reflex EMG that we observed. Indeed, a two-way ANOVA showed that, for the group, there was no significant variation in background EMG throughout the cycle (P > 0.99).


    DISCUSSION
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES

This study extends the results of previous investigations on the effects of imposed movement on reflexes in two important ways. First, by using a precise mechanical stimulus to elicit the stretch reflex, rather than an electrical stimulus to evoke the H-reflex, it assessed the excitability of the complete reflex circuit including the muscle receptors. Second, by measuring the reflex torque as well as the EMG, it evaluated the mechanical consequences of the reflex response.

Comparison with previous results

Imposed movements of either the hip or the knee have been shown to inhibit the soleus H reflex (Brooke et al. 1993, 1995a,b; Cheng et al. 1995a,b; Collins et al. 1993; Hultborn et al. 1996; Nielsen et al. 1993; Voigt and Sinkjaer 1998). In some experiments, simulating stepping, the effects of passive ankle movement were small (Brooke et al. 1995a). However, in other experiments substantial decreases in H reflexes were observed after passive and active movements of the ankle (Hultborn et al. 1996; Nielsen et al. 1992) and during imposed sinusoidal movements (Voigt and Sinkjaer 1998). In the present experiments, the stretch reflex response dropped significantly throughout the simulated walking cycle in all subjects. The different results may be due, in part, to differences in the range of motion and angular velocity used in the various experiments.

The modulation of the soleus stretch reflex EMG response observed in this study was strikingly similar to that reported recently for normal (active) walking (Andersen and Sinkjaer 1995). It also closely resembles the modulation of H-reflex gain during walking (Capaday and Stein 1987). Indeed, Andersen and Sinkjaer (1995) reported no significant difference in the modulation of stretch reflex and H-reflex responses throughout most of the walking cycle except in late stance where the H reflex was significantly larger.

There are, however, some differences between our results and those reported by Sinkjaer et al. (1996). First, these researchers reported a recovery of the stretch reflex late in the swing phase of normal walking. We saw no such recovery; the reflex response remained low throughout the swing phase. Second, they found no significant difference between the amplitude of the stretch reflex during the stance phase of walking and under static conditions at matched soleus EMG activity. In our study, the reflex EMG clearly was inhibited during movement in comparison with matched static conditions. We do not fully understand these differences, but there were a number of methodological differences (see Other factors). Further studies will be needed to test which of these variables accounts for the different findings.

Reflex torque

A number of previous studies have documented reflex torques generated under steady-state conditions (Carter et al. 1990; Crago et al. 1977; Nichols and Houk 1976; Sinkjaer and Hoffer 1990; Sinkjaer et al. 1988). However, studies of reflex modulation during movement have focused on EMG. We believe that this study is the first to measure corresponding reflex torque in humans. One important finding in this study is that during movement, the amplitude of the reflex EMG does not always provide a good measure of the torque that is generated. Figure 4 illustrates this very clearly; the pulse that elicited the greatest EMG response gave rise to little or no reflex torque. This dissociation between reflex EMG and reflex torque implies that muscle dynamics (i.e., the force-length and -velocity properties of muscle) play a critical role in determining the functional significance of the stretch reflex during movement.

To investigate the effects of muscle dynamics on reflex torque, we plotted the ratio of reflex torque to reflex EMG as a function of position and velocity (Fig. 7). Under static conditions (), where joint velocity is zero, the ratio was at a minimum near midposition and increased monotonically as the ankle was moved toward its limits, consistent with the force-length properties of muscle. During imposed walking (open circle ), the reflex torque/EMG ratio was low at long muscle lengths and high shortening velocities. Conversely it was high when the muscle was short and being stretched. Unfortunately it is impossible to dissociate the effects of position and velocity because these variables covaried in our data set. The reversal of the dependence on position from static conditions to imposed walking does suggest, however, that force-velocity effects dominate during this movement.



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Fig. 7. Ratio of reflex torque to reflex EMG for static () and imposed walking trials (open circle ) as a function of position and "walking" velocity. Time relative to heel strike at which the pulse was applied also is indicated for the imposed walking data. All data are group means. Points where the reflex EMG was negligible (<5% of the maximum) are not shown; consequently the range of velocities is less than in Fig. 6.

A dissociation in changes in reflex torque and reflex EMG has been observed under static conditions as a function of activation level (Toft et al. 1991). This was attributed to the effects of increased motor neuron synchronization. Such mechanisms may contribute to the results of the present experiment but cannot explain our observation that there was little or no reflex torque associated with the largest reflex EMG.

To what extent do the responses to sudden stretches provide insight into the role of the stretch reflex during normal walking or in response to natural perturbations? In both cases, the position changes are likely to be much larger and the velocities much lower than for the pulses. Certainly, in view of the nonlinear dependence of the stretch reflex on muscle velocity and activation (Stein and Kearney 1995), we cannot predict the reflex torques generated under "natural" conditions from the responses to rapid stretches. However, the high sensitivity of the stretch reflex to the pulse perturbations makes it unlikely that slower, natural movements would evoke a response when pulses failed to. This suggests that the stretch reflex might generate substantial torques during the earlier parts of the step cycle both during normal walking, where the muscle is being stretched under the weight of the body, and in response to naturally occurring perturbations. In contrast, our results suggest that the stretch reflex would contribute little torque during the push-off and later phases.

Methodological considerations

The major conclusion of this study is that peripheral input from the ankle contributes strongly to the cyclic modulation of the stretch reflex in the soleus muscle. For this conclusion to be valid, we must rule out other possible sources of the modulation. One possibility is that the reflex modulation results from the cyclic modulation of voluntary drive locked to the ankle movement. We attempted to eliminate this possibility experimentally by asking subjects to generate a constant level of plantarflexing torque and heavily low-pass filtering the feedback signal to eliminate any cyclic clues. These attempts were not completely successful but any variations that remained in the behavior of individual subjects were not correlated with the systematic reflex responses that were seen in all subjects.

Another possible source of modulation could be systematic variations in the pulse properties. We designed the experiments to maintain the incremental pulse velocity constant and, as illustrated in Fig. 6A, this was achieved very well. The net velocity (the sum of velocity of the pulse and of the underlying walking pattern) did vary somewhat throughout the cycle. However, changes in the reflex amplitude were not correlated with the net stretch velocity of the muscle (or, more precisely, with the net angular velocity of the joint when a pulse was applied). We conclude therefore that the modulation of the reflex during the movement was not due to differences in the net velocity.

Other factors

There are clearly a number of important differences between the imposed walking task and normal walking that must be discussed to assess the significance of our findings. These include the following: 1) the range of soleus activation is much larger during normal gait than the small changes observed in the course of this experiment; 2) the subjects were lying down during the experiment and so experienced different vestibular and cutaneous inputs than during normally walking; and 3) walking involves simultaneous movement of many joints.

A constant (or relatively constant) descending drive to the triceps surae was necessary in our study to explore the role of peripheral afferents in the modulation of the stretch reflex. Because reflex amplitude depends on the excitation level of the motoneuron pool, the present experimental procedures should be applied in the future at various levels of voluntary contraction to ensure that the role of peripheral afferents observed in this study holds for the range of excitation levels seen during walking. Fatigue of the relevant muscles may limit the extent to which this is possible.

Positioning the body horizontally rather than vertically has been reported to somewhat increase the amplitude of the H reflex (Brooke et al. 1995a; Misiaszek et al. 1995). Nevertheless the modulation pattern observed in this study closely resembled that reported for walking, suggesting that vestibular and cutaneous inputs play relatively minor roles in the modulation of the stretch reflex during this movement.

By imposing movement about the ankle only, we were able to isolate the effect to afferents associated with this joint and with muscles crossing this joint. Moving other joints, such as the hip or the knee, of either the ipsilateral or the contralateral leg, inhibits the soleus H reflex (Brooke et al. 1993; Collins et al. 1993). Such influences may contribute to the state of the soleus stretch reflex gain during normal walking. However, the similarity between the stretch reflex modulation during imposed walking movement of the ankle and normal walking suggests that afferents associated with the ankle joint or with muscles crossing this joint play a major role in reflex modulation during normal walking.

Thus although various considerations led us to conduct experiments that differ in several important respects from normal walking, we are confident that our results have important implications for walking.

Possible neural mechanisms

What afferents are responsible for the reflex modulation seen in the present study? Previous work, on the effect of passive movement of the knee on the H reflex from a distal plantar muscle in dogs and on the soleus H reflex in humans, points to spindle afferents from the knee extensor muscles as an important source of soleus reflex depression (Cheng et al. 1995b; Misiaszek et al. 1995). The strong relation between the mean absolute velocity of movement and reflex depression described in our previous paper (Stein and Kearney 1995) is certainly consistent with the hypothesis that spindle afferents are involved, as is the observation that presynaptic inhibition has tonic and phasic components (Stein 1995). Mechanisms other than presynaptic inhibition may be involved such as the homosynaptic postactivation depression described by Hultborn (Hultborn et al. 1996). However, as argued in Voigt and Sinkjaer (1998) it is difficult to see how these could account for the modulation of the reflex throughout the walking cycle.

Conclusion

The pattern of stretch reflex modulation observed during imposed walking movement of the ankle joint is very similar to that of the stretch and H reflexes during real walking. This suggests that during walking peripheral afferents from ankle muscles play an important role in modulating the sensitivity of the stretch reflex to both sudden stretches and natural inputs. Furthermore the dissociation between reflex EMG and reflex torque observed in this study demonstrates that muscle mechanics play a key role in determining the functional importance of the stretch reflex during movement.


    ACKNOWLEDGMENTS

This work was supported by grants from the Medical Research Council of Canada and the Natural Sciences and Engineering Research Council of Canada.


    FOOTNOTES

Address for reprint requests: R. E. Kearney, Dept. of Biomedical Engineering, McGill University, 3775 University St., Montreal, Quebec H3A 2B4, Canada.

The costs of publication of this article were defrayed in part by the payment of page charges. The article must therefore be hereby marked "advertisement" in accordance with 18 U.S.C. Section 1734 solely to indicate this fact.

Received 11 December 1998; accepted in final form 1 March 1999.


    REFERENCES
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ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
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0022-3077/99 $5.00 Copyright © 1999 The American Physiological Society