1Department of Biomedical Engineering, McGill University, Montreal, Quebec H3A 2B4; and 2Division of Neuroscience, University of Alberta, Edmonton, Alberta T6G 2S2, Canada
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ABSTRACT |
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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.
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INTRODUCTION |
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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
-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
).
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METHODS |
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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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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.
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RESULTS |
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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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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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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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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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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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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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The velocity of the imposed walking movements varied from 2 to +1
rad/s as shown in Fig. 6A (
). Pulses were
superimposed on these movements so that the net velocity, the sum of
the walking and pulse velocities (
), 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 () in Fig. 6B had two peaks of EMG, one in
midstance and one in swing. Another subject (
) 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).
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DISCUSSION |
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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 (
), 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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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.
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ACKNOWLEDGMENTS |
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This work was supported by grants from the Medical Research Council of Canada and the Natural Sciences and Engineering Research Council of Canada.
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FOOTNOTES |
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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.
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REFERENCES |
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