1Rehabilitation Research and Development Center (153), Veterans Affairs Palo Alto Health Care System, Palo Alto 94304-1200; and 2Mechanical Engineering Department (Biomechanical Engineering Division) and 3Department of Functional Restoration, Stanford University, Stanford, California 94305-3030
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ABSTRACT |
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Ting, Lena H., Steven A. Kautz, David A. Brown, and Felix E. Zajac. Contralateral Movement and Extensor Force Generation Alter Flexion Phase Muscle Coordination in Pedaling. J. Neurophysiol. 83: 3351-3365, 2000. The importance of bilateral sensorimotor signals in coordination of locomotion has been demonstrated in animals but is difficult to ascertain in humans due to confounding effects of mechanical transmission of forces between the legs (i.e., mechanical interleg coupling). In a previous pedaling study, by eliminating mechanical interleg coupling, we showed that muscle coordination of a unipedal task can be shaped by interlimb sensorimotor pathways. Interlimb neural pathways were shown to alter pedaling coordination as subjects pedaling unilaterally exhibited increased flexion-phase muscle activity compared with bilateral pedaling even though the task mechanics performed by the pedaling leg(s) in the unilateral and bilateral pedaling tasks were identical. To further examine the relationship between contralateral sensorimotor state and ipsilateral flexion-phase muscle coordination during pedaling, subjects in this study pedaled with one leg while the contralateral leg either generated an extensor force or relaxed as a servomotor either held that leg stationary or moved it in antiphase with the pedaling leg. In the presence of contralateral extensor force generation, muscle activity in the pedaling leg during limb flexion was reduced. Integrated electromyographic activity of the pedaling-leg hamstring muscles (biceps femoris and semimembranosus) during flexion decreased by 25-30%, regardless of either the amplitude of force generated by the nonpedaling leg or whether the leg was stationary or moving. In contrast, rectus femoris and tibialis anterior activity during flexion decreased only when the contralateral leg generated high rhythmic force concomitant with leg movement. The results are consistent with a contralateral feedforward mechanism triggering flexion-phase hamstrings activity and a contralateral feedback mechanism modulating rectus femoris and tibialis anterior activity during flexion. Because only muscles that contribute to flexion as a secondary function were observed, it is impossible to know whether the modulatory effect also acts on primary, unifunctional, limb flexors or is specific to multifunctional muscles contributing to flexion. The influence of contralateral extensor-phase sensorimotor signals on ipsilateral flexion may reflect bilateral coupling of gain control mechanisms. More generally, these interlimb neural mechanisms may coordinate activity between muscles that perform antagonistic functions on opposite sides of the body. Because pedaling and walking share biomechanical and neuronal control features, these mechanisms may be operational in walking as well as pedaling.
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INTRODUCTION |
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Recent findings about the ability to improve the
locomotor capacity of individuals with spinal cord injury through the
application of principles derived from pattern generator theory (for
review, see Rossignol et al. 1996) underscore the
necessity to better understand interlimb coordination mechanisms in
humans; this would lead to the development of more effective
rehabilitation strategies. The human spinal cord may possess some
ability to produce bilateral locomotor activity (Calancie et al.
1994
; Harkema et al. 1997
; Rossignol et
al. 1996
) as shown in a number of vertebrates (for review, see
Rossignol 1996
). How such centrally generated signals are integrated with peripheral afferent information to produce coordinated bilateral locomotion remains elusive, especially in humans.
However, motor patterns produced by individuals with spinal cord injury
depend on bilateral sensory information associated with limb movement
and loading (Harkema et al. 1997
).
In vertebrate preparations, no clear picture of how sensory information
modulates ipsi- and contralateral patterns exists. Evidence from spinal
cats show that locomotor activity in each hindlimb can be generated
independently (e.g., Grillner and Zangger 1979).
Ipsilateral sensory feedback has been shown to be important in
reinforcing locomotor activity, and both ipsilateral hip angle and
loading of extensors affect ipsilateral phase changes from stance to
swing in cat (Duysens and Pearson 1980
; Grillner
and Rossignol 1978
; Pearson et al. 1992
). In
split-belt treadmill conditions, independent rhythm generation in each
hindlimb is demonstrated by the ability of the hindlimbs to walk at
different speeds (Forssberg et al. 1980
). Further, one
hindlimb can continue rhythmic behavior even when the other is
prevented from doing so (Duysens and Pearson 1980
;
Grillner and Rossignol 1978
). On the other hand, in
similar cat preparations, interdependence of sensorimotor signals in
the two hindlimbs is demonstrated by the maintenance of integral (e.g.,
1:1 or 1:2) step frequencies, which ensures that one foot is always on
the ground (Forssberg et al. 1980
). Following unilateral
deafferentation in spinal cats, disruption of both ipsi- and
contralateral stepping occurs; this further illustrates the
contralateral influence of afferent input (Giuliani and Smith
1987
). In turtles, bilateral coupling of centrally generated rhythmic output is also demonstrated because spinal cord hemisection alters bilateral fictive rhythmic activity (Stein et al.
1995
).
Elucidation of interlimb neural coupling mechanisms in humans is even
more challenging because central and peripheral influences cannot be
explicitly isolated. Changes in muscle coordination of a leg in
unilateral tasks compared with similar bilateral tasks may be caused by
two major factors: 1) contralateral sensorimotor signals
mediated through neural interlimb coupling mechanisms and 2)
ipsilateral afferent signals triggered by the forces transmitted to the
ipsilateral leg due to the acceleration or movement of the
contralateral leg. Although clear evidence for neural interlimb coupling has been demonstrated in static tasks in which no loadsharing occurs between the limbs (Howard and Enoka 1991;
Schantz et al. 1989
; Secher et al. 1988
),
results from dynamic tasks are not as conclusive. For example, though
perturbations in stance elicit bilateral electromyographic (EMG)
responses of similar latencies (Berger et al. 1984
;
Dietz and Berger 1982
), the EMG changes in the
nonperturbed leg may be due to afferent signals generated in that leg
as a consequence of the instantaneous joint reaction forces generated
in both legs, and the subsequent motion of the limb segments in the
nonperturbed leg (Yamaguchi and Zajac 1990
; Zajac
1993
). Therefore mechanical and neuronal coupling can
simultaneously affect muscle coordination of movement, making it
difficult to isolate the effects of either factor. Nevertheless it
seems reasonable that coordinative neuronal coupling would be effective
in recovery from tripping or other perturbations to normal gait (e.g.,
Berger et al. 1984
; Dietz et al. 1986
;
Eng et al. 1994
).
Pedaling is a useful paradigm in the study of human locomotion. Task
mechanics can be controlled and manipulated. The alternating flexion
and extension of the limbs, characteristic of many modes of locomotion,
can be studied without the confounding influence of balance. Phasing
and frequency of leg movements are similar in walking and pedaling.
Further, because the subject is seated, neither balance nor body-weight
support is required, and the kinetics and kinematics of the legs can be
analyzed in isolation of the head, arms, and trunk. In both pedaling
and walking, significant forces are generated by the legs and applied
to the environment during the extension phase, and passive or external
forces tend to flex the limb during the flexion phase (Eng et
al. 1997; Kautz and Hull 1993
). In addition,
many reflexes such as the H reflex and flexor reflex are modulated over
the gait cycle and similarly modulated over the pedaling cycle (e.g.,
Brown and Kukulka 1993
; for review, see Brooke et
al. 1997
). Thus neuronal elements responsible for gait may
participate in pedaling tasks as well.
This study focused on neural interlimb coupling mechanisms that operate
during the flexion phase in steady-state pedaling because a previous
pedaling study (Ting et al. 1998b) found that the most
significant difference between bilateral and unilateral pedaling occurs
in the flexion phase. Ting et al. proposed that sensorimotor signals
associated with contralateral extension play a role in modulating
flexion-phase muscle activity in pedaling. Because force and movement
may provide powerful influences on the ongoing locomotor pattern, the
effects of contralateral movement and extensor force production on
muscle activity in the flexion phase during unilateral pedaling were
investigated. The sensorimotor conditions in the contralateral
nonpedaling leg were designed to mimic pedaling in the amount and
timing of the generation of extensor force, the largest component of
muscular force production during pedaling (Kautz and Hull
1993
; Raasch et al. 1997
), and/or in the
antiphasing of the movement of the legs. Specifically, we hypothesized
that isolated generation of forces and/or passive movements of the
contralateral nonpedaling leg would inhibit the flexion phase activity
normally present in a unilateral pedaling task. Abstracts of this work
have appeared (Ting et al. 1997
, 1998a
).
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METHODS |
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Eighteen healthy subjects [10 male, 8 female; age, 22 ± 3 (SD) years; height, 1.7 ± 0.7 (SD) m; weight, 65 ± 8 (SD) kg] who rode a bicycle for <50 miles/wk and who were naive to the experimental goals signed consent forms prior to participation in the study. This study was approved by the Institutional Review Board (Medical Committee for the Protection of Human Subjects in Research) at Stanford University Medical School.
The pedaling leg (left leg) of each subject performed the same pedaling task in all trials, while the condition of the nonpedaling (right leg) leg was varied. The conditions of the nonpedaling leg were chosen such that the effects of leg movement and extensor force generation in that leg on flexion-phase coordination of the pedaling leg could be tested. The interaction effects of nonpedaling leg force and movement were also tested along with the effect of force level (i.e., low vs. high relative to typical forces encountered during pedaling) and rhythmicity (i.e., rhythmic application of force).
Experimental apparatus
A bicycle ergometer was modified so subjects would pedal against the same mechanical load profile for all eight conditions. The left and right cranks were mechanically uncoupled. Thus the right leg could not mechanically influence left leg pedaling coordination. In addition, the kinematic relationship between the left and right cranks could be manipulated using a servomotor. Subjects were linked to the cranks via clipless pedals and standard bicycling cleats; thus the feet maintained contact with the pedals at all times.
The left crank propelled a flywheel, which was removed from a Monark bicycle ergometer. A constant-force spring (or a negator spring, i.e., a spring that develops the same force irrespective of its length), attached to the left crank arm (Fig. 1A), provided an approximately sinusoidal crank torque that impeded propulsion during extension (0-180°, Fig. 3A) and aided propulsion during flexion (180-360°). This sinusoidal load pattern approximately replicated the alternating pattern of retarding and propulsive torque generated by the right leg in two-legged pedaling (Fig. 1A). The peak crank torque contributed by the spring (22 N-m) was chosen such that the leg would be propelled to overcome gravity in the flexion phase. Thus as in two-legged pedaling, subjects could have executed the pedaling task with little active flexion.
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The nonpedaling (right) crank was controlled by a programmable servomotor (Kollmorgen B606A motor, D20 motor controller, 2 kHz servo loop; Kollmorgen Motion Technologies Group, Commack, NY), which either fixed the nonpedaling crank in a static position or moved the crank at a 180° phase relation to the opposite (left) crank (Fig. 1B). In the movement condition, the motor was servocontrolled by the optical encoder signal from the opposite (left) pedaling-leg crank. As a result, torque applied by the nonpedaling leg did not contribute to crank propulsion because all torque generated by the nonpedaling leg was resisted by the motor as it maintained the 180° phase relation with the pedaling crank. The motor was also used in a control condition where it rotated the crank at a constant 60 rpm.
Subjects wore a cleated cycling shoe on the pedaling foot and a cleated
ankle brace (DePuy Orthotech, Tracy, CA) fixed at 10° plantarflexion
on the nonpedaling leg (Fig. 1B). By fixing the ankle angle,
the configuration of the leg (hip and knee angles) is uniquely
determined by the crank angle (Fregly and Zajac 1996; Kautz and Hull 1993
; Redfield and Hull
1986
). Thus no muscular effort was required to maintain a
static configuration when the crank position was fixed on the
nonpedaling side. Further, no muscular effort was required to maintain
a comfortable limb trajectory when the crank was moved by the motor.
Because the relationship between contralateral extension and
ipsilateral flexion was being explored, the ankle angle of the brace
was set to 10° plantarflexion to correspond to the position the ankle
has in the mid-extension phase during pedaling (~107°, Fig.
3A) (Nordeen-Snyder 1977
). Subjects were
seated during all trials and wore a hip belt to reduce pelvic motion.
Subjects sat with a forward lean of ~10°, which is a typical trunk
angle during walking (Pozzo et al. 1990
), and supported
some of their torso weight on the handlebars (Ting et al.
1998b
).
Pedal forces were measured using pedal dynamometers on both the
pedaling and nonpedaling crank (Newmiller et al. 1988).
A load cell measured the force of the constant force spring. Crank and
pedal angles from both sides were measured using digital optical encoders (4096 counts/revolution). A linear transducer was attached to
the hip belt over the anterior superior iliac spine (ASIS) of the
nonpedaling leg, which measured the amplitude of ASIS movement in it's
primary direction of movement which is forward, downward and laterally
(cf. Neptune and Hull 1995
). This measure of ASIS movement was used to compare pelvis motion between different trial conditions.
Surface EMGs were measured from seven muscles bilaterally: vastus medialis (VM), left (LVM) and right (RVM); rectus femoris (RF), left (LRF) and right (RRF); biceps femoris long head (BF), left (LBF) and right (RBF); semimembranosus (SM), left (LSM) and right (RSM); tibialis anterior (TA), left (LTA) and right (RTA); medial gastrocnemius (MG), left (LMG) and right (RMG); and soleus (SL), left (LSL) and right (RSL).
All signals were sampled at 1,000 Hz. Analog RC anti-aliasing filters with a cutoff frequency of 80 and 800 Hz were used on non-EMG and EMG channels, respectively, to reduce very high-frequency noise from the motor (~20 kHz).
Practice protocol
Subjects were trained to pedal with their pedaling leg so that
they could maintain a constant cadence of 60 rpm without using any
feedback. A preliminary trial oriented subjects to the apparatus and
ensured that seat and handlebar heights were appropriate. Subjects
pedaled bilaterally with the motor driving the right crank antiphase to
the left pedaling-crank encoder signal. Subjects perceive this
situation as normal two-legged pedaling. A metronome helped subjects
maintain a steady cadence during the first 40 s of the trial.
Next, five practice trials of 60-s duration of unilateral pedaling were
presented to subjects. The metronome usage duration was decreased
incrementally from 40 to 10 s across the five trials. To avoid
cumulative fatigue, subjects rested for 1 min between each practice trial.
After the practice session, subjects were able to pedal smoothly and
consistently at a steady cadence between 55 and 65 rpm. Smoothness was
determined by the absence of freewheeling, which is a decoupling of the
crank from the flywheel load that occurs when the crank decelerates
relative to the flywheel (for discussion, see Fregly and Zajac
1996; Raasch 1995
). In a few cases, subjects did
not perform consistently after five practice trials and were given
additional practice trials.
Next, to familiarize subjects with the sensation of having their right leg moved by a motorized crank, the nonpedaling crank was rotated by the motor at 60 rpm so subjects would be able to relax their right leg in subsequent movement conditions when requested. During this practice trial the pedaling leg was relaxed with the pedaling crank fixed in a horizontal, mid-extension position (~107° crank angle).
Finally, subjects practiced using the vertical force-feedback indicator employed during the experimental conditions. Both pedaling and nonpedaling cranks were locked at mid-extension (~107° crank angle). The level of downward force being exerted perpendicular to the surface of the pedal of the nonpedaling leg was displayed on a computer monitor, which displayed a vertical bar. The vertical bar was divided into five regions. Each bar was illuminated when the force reached the weight of the leg resting on the pedal at mid-extension plus a specified amount. Two of the bars were larger in height to indicate a larger range of force. These were the "target" zones of the high- and low-force levels. The three smaller bars served to indicate when the subject exceeded or fell short of the targets. The high-force level corresponded to the force on the pedal during mid-extension of pedaling [~300 N; i.e., (leg weight +150 N) to (leg weight + 450 N)]. The low level was just slightly above the weight of the leg [i.e., (leg weight + 25 N) to (leg weight + 100 N)]. The height of the force windows were chosen such that changes in force due to subtle movement or shifting of position of the right leg would not cause the subject to exit the desired force range. Further, it was important that subjects could generate the force easily and naturally, as in normal pedaling, exerting as little cognitive control over force level as possible. Four 20-s practice trials were performed. Subjects were asked to maintain a tonic level of extensor force in the high and then low target range for the duration of the trial. Then they were asked to generate rhythmic extensor force (paced by a 1-Hz metronome), first to the high target and then to the low target.
Experimental conditions
Eight trial conditions and one additional control condition were presented to subjects in random order. In all but the control condition, the pedaling leg pedaled at 60 rpm. Each trial condition was given a two-letter code indicating the state of the nonpedaling leg (Fig. 2): either static (S) or moving (M) and either generating a force (F) or relaxed (R). A subscript was used to indicate tonic (T) or rhythmic (R) force with a capital letter to designate a high level of force and a lower-case letter to designate a low level of force. In the cases where the nonpedaling leg was static, the nonpedaling crank was maintained in the horizontal position (107°), which corresponds approximately to mid-extension in pedaling. In the movement conditions, the nonpedaling crank was servomoved antiphase to the left crank. In the cases where the nonpedaling leg was relaxed and did not generate force, the subject was instructed to neither resist nor aid the motion imposed by the motor. In force-generating conditions, subjects used the force feedback indicator to achieve the desired force level.
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Four trial conditions were used to investigate the importance of contralateral nonpedaling extensor force generation and movement on flexion-phase coordination of the ipsilateral pedaling leg (Fig. 2A). The nonpedaling leg was either static and relaxed (SR; i.e., the nominal condition), moving and relaxed (MR), static and producing a high rhythmic isometric extensor force (SFR), or moving and producing a high rhythmic extensor force (MFR). During rhythmic force production, subjects were instructed to push downward with the nonpedaling leg concurrent to the flexion phase of the pedaling leg, such that the timing would be similar to that of two-legged pedaling. Thus subjects were asked to generate an extensor force to the high-force target.
Two additional trial conditions were used to determine the effect of extensor force level produced by the nonpedaling leg. Subjects generated a rhythmic force with the contralateral nonpedaling leg static or moving, as before, but only to the low-force target (SFr, MFr).
Finally, two other trial conditions were used so that the effect of the
rhythmic nature of force generation compared with tonic force
generation could be studied. In these two conditions, subjects
maintained a constant force level with the nonpedaling leg throughout
the pedaling cycle (SFt,
SFT, Fig. 2C). This was only done in
the static condition because additional constraint and
movement-dependent forces on the pedal (Kautz and Hull
1993) make it impractical to measure constant force production
by the subjects when the leg is moving. Further, subjects would have had to execute the difficult task of resisting the motor while simultaneously allowing the motor to move the nonpedaling leg.
In the "control" condition (MRØ), the usually pedaling leg was stationary and relaxed while the nonpedaling leg, also relaxed, was moved at 60 rpm by the motor for ~40 s. Notice that this "control" condition, where neither leg pedals, is different from the nominal trial condition (SR).
Each trial condition was also ~40 s in duration, with the metronome used in the first 10 s, and data collected in the last 20 s of the trial after the subjects had reached a steady-state cadence without the metronome. To minimize the effects of fatigue, subjects rested for at least one minute between trials.
Data processing
Force and angle data were downsampled to 200 Hz, low-passed filtered using a Butterworth filter (10-Hz cutoff, zero-lag), and used to calculate crank torque. Crank torque is the component of pedal force, multiplied by crank arm length, which accelerates the crank. All data were referenced to pedaling-leg crank angle, with 0° corresponding to the crank being closest to the seat (Fig. 3A). Crank angles between 0 and 180° refer to periods of leg extension, when the foot is moving away from the pelvis. Crank angles between 180 and 360° refer to leg flexion, when the foot moves toward the pelvis. Data from each condition of each subject were averaged over 10 consecutive crank revolutions.
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EMG signals were high-pass filtered with a Butterworth filter (35-Hz cutoff, 0 lag) to remove low-frequency motor noise. Mean signal offset was subtracted from the EMG signals and EMGs were rectified before further processing.
For data analysis, the signals measured during the pedaling crank cycle
were divided into four quadrants (Fig. 3A) with each centered at mid-phase of one or two of the six biomechanical functions executed in forward pedaling (Fig. 3B) (Raasch and
Zajac 1999; Raasch et al. 1997
; Ting et
al. 1999
). Thus flexion-phase coordination of pedaling was
quantified by analyzing data in quadrant 4, and changes in the
extension-to-flexion phase were quantified by analyzing data in
quadrant 3, etc. Data analysis focused on flexion-phase (quadrant 4) coordination.
Pedaling-leg integrated EMG (iEMG) and work output in each quadrant were calculated and averaged over all steady-state cycles (~20) to produce mean values for each subject and trial condition. The net work output by the pedaling leg in each quadrant is proportional to the average crank torque in the quadrant. Work output (workload) and iEMG over the entire cycle were found by summing the respective quantities over the four quadrants.
Performance of the nonpedaling leg was also monitored in each trial to
ensure that the desired sensorimotor state was achieved. Mean extensor
force generated by the nonpedaling leg concurrent with the flexion
phase of the pedaling leg (quadrant 4) was calculated, along with iEMG
from the nonpedaling leg and the standard deviation of ASIS motion on
the nonpedaling side. Because EMG signals during static and dynamic
conditions cannot be directly compared due to differences in motor unit
and muscle recruitment (Gielen 1999; Theeuwen et
al. 1994
; van Bolhuis and Gielen 1997
),
comparison of nonpedaling EMGs among only static conditions, or among
only movement conditions, were made.
Data analysis
Data were analyzed (2-way ANOVA with subject and trial condition as a factor) to answer the following questions:
1) Does nonpedaling leg movement and/or force generation affect flexion-phase coordination in the pedaling leg? [2-way ANOVA comparing only trial conditions SR, MR, SFR, and MFR to test independent and combined effects of movement and high rhythmic force generation (MR, SFR, and MFR) relative to the nominal condition (SR)].
2) Does the level of force generated by the nonpedaling leg affect flexion-phase coordination in the pedaling leg? (2-way ANOVA with trial conditions SFR, MFR, SFr, and MFr to test effects of high vs. low rhythmic force, with and without movement; and with trial conditions SR, MR, SFr, and MFr to test the effects of low rhythmic force vs. no force generation.)
3) Is the presence of force generation alone during nonpedaling leg extension sufficient to affect pedaling flexion-phase coordination or is the additional rhythmic nature of the force also contributory? Specifically, does tonic force generation have the same effect as rhythmic force generation? (2-way ANOVA with SFR, SFr, SFT, and SFt to test the effects of tonic vs. rhythmic force generation at high and low levels.)
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RESULTS |
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Force data from the pedaling leg were only available for 14 of the 18 subjects due to a damaged wire. All other data, including pedaling-leg EMGs, were collected from all 18 subjects.
Performance of subject and experimental apparatus
The workload and cadence of the pedaling leg remained unchanged
over all conditions in all subjects. The average workload for the
subjects ranged from 79.7 to 82.1 J/cycle and was not significantly
different for any condition (P > 0.05 for all pairwise comparisons). The work done by the constant force spring was also consistent across all conditions (P > 0.05 for all
pairwise comparisons) and varied by <1% during any particular trial.
Subjects were able to maintain a pedal cadence of ~60 rpm in all
trials, with mean cadences per subject ranging from 59 to 67 rpm. ASIS
motion was cyclical, generally moving forward and downward during the
downstroke (cf. Neptune and Hull 1995) with peak forward
displacement occurring when the crank was near bottom-dead-center
(example: Fig. 4, A and
B). Although hip movement was greater in movement conditions than in static conditions, there was no difference in amplitude of hip
motion of the nonpedaling leg across all static trials or across all
movement trials (Fig. 4C).
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The servomotor was successful in achieving the desired isometric and movement conditions of the nonpedaling crank. During static conditions, the nonpedaling crank position was maintained in a horizontal position [0.3 ± 0.4° (SD) average position error across all subjects]. The crank did not move appreciably as the standard deviation of movement within each trial was near zero (0.07 ± 0.05° across all trials). In movement conditions, the desired 180° antiphase relationship was essentially maintained, as the average phase between left and right cranks across all trials was 176 ± 5°, with an average standard deviation of 1.4 ± 1.1° within each trial.
Effects of nonpedaling leg movement and force generation on pedaling coordination
The condition where the nonpedaling leg was static and relaxed
(SR) served as the nominal condition against which the movement and
force generating conditions of the nonpedaling leg were compared. The
flexion-phase torque profile and EMGs of the pedaling leg were similar
to those reported previously for unilateral pedaling where the
contralateral nonpedaling leg was in a similar sensorimotor state
(Ting et al. 1998b). Crank torque in the flexion phase
(quadrant 4) was near zero (example: Fig.
5A), and correspondingly,
flexion-phase work output across all subjects was also about zero (Fig.
6A), which differs from the
negative work output normally observed in two-legged pedaling (about
10 J) (Ting et al. 1998b
).
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Compared with the nominal condition (SR), passive movement (MR) of the nonpedaling leg was unsuccessful in reducing the crank torque and work output produced by the pedaling leg during its flexion phase [crank torque example from 1 subject: compare Fig. 5B (quadrant 4) with Fig. 5A (quadrant 4); work output across all subjects: Fig. 6A, compare MR with SR, P > 0.05]. Correspondingly, EMG activity in the pedaling leg during the flexion phase was also unchanged [example from 1 subject: compare Fig. 7B (quadrant 4) with Fig. 7A (quadrant 4); average iEMG across all subjects: Fig. 8, A-D, compare MR with SR, all P > 0.05].
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In contrast, high rhythmic extensor force generation in the stationary nonpedaling leg (SFR) was successful in reducing flexion-phase crank torque and work output in the pedaling leg compared with the nominal condition SR [crank torque example from 1 subject: compare Fig. 5C (quadrant 4) with Fig. 5A (quadrant 4); work output across all subjects: Fig. 6B, compare SFR with SR, P < 0.01]. Correspondingly, EMG activity over the flexion phase decreased in some of the pedaling leg muscles, specifically in BF and SM [example from 1 subject, compare LBF and LSM in Fig. 7C (quadrant 4) with LBF and LSM in Fig. 7A (quadrant 4); average iEMG across all subjects: Fig. 8, A and B, compare SFR with SR, both P < 0.01]. However, no reduction of iEMG activity in RF or TA over the flexion phase were observed [example from 1 subject, compare LRF and LTA in Fig. 7C (quadrant 4) with LRF and LTA in Fig. 7A (quadrant 4); average iEMG across all subjects: Fig. 8, C and D, compare SFR with SR, both P > 0.05].
The addition of movement to high rhythmic force generation in the nonpedaling leg (MFR) was sufficient, however, to reduce RF and TA activity in the pedaling leg during its flexion phase [example from 1 subject, compare LRF and LTA in Fig. 7D (quadrant 4) with LRF and LTA in Fig. 7A (quadrant 4); average iEMG across all subjects: Fig. 8, C and D, compare MFR with SR, P < 0.05 for LRF and P < 0.01 for LTA]. This was the only condition where RF and TA activity were reduced (Fig. 8, C and D: compare MFR with the other conditions). This condition (MFR) also reduced BF and SM activity in the pedaling leg [example from 1 subject, compare LBF and LSM in Fig. 7D (quadrant 4) with LBF and LSM in Fig. 7A (quadrant 4); average iEMG across all subjects: Fig. 8, A and B, compare MFR with SR, both P < 0.01]. However, the addition of movement to the nonpedaling leg did not reduce further the reduction in BF and SM activity observed with high rhythmic force generation alone (Fig. 8, A and B, compare MFR with SFR, both P > 0.05). Similarly, flexion-phase work output during this condition of movement and high rhythmic force generation in the nonpedaling leg (MFR) decreased compared with the nominal condition [crank torque example from 1 subject: compare Fig. 5D (quadrant 4) with Fig. 5A (quadrant 4); work output across all subjects: Fig. 6B, compare MFR with SR, P < 0.01] but no more than the decrease observed with high rhythmic force generation (Fig. 6B, compare MFR with SFR, P > 0.05).
A low level of rhythmic force generation by the nonpedaling leg, whether the leg was moved (MFr condition) or not (SFr condition), was also effective in reducing both the flexion-phase work output in the pedaling leg (compare MFr and SFr in Fig. 6C with SR in Fig. 6A; both P < 0.01) and the activity in BF and SM (Fig. 8, A and B, compare MFr and SFr with SR; both P < 0.01) but was ineffective in reducing RF and TA activity (Fig. 8, C and D, compare MFr and SFr with SR; both P > 0.05). This reduction in BF and SM activity when rhythmic force generation was low did not differ from the decrease found when rhythmic force generation was high (Fig. 8, A and B, compare MFr and SFr with MFR and SFR; all P > 0.05). However, the reduction in flexion-phase work output in the static rhythmic low-force condition (SFr) was less than the reduction in either the rhythmic high-force static condition (compare SFr in Fig. 6C with SFR in Fig. 6B; P < 0.05) or the rhythmic high-force moving-leg condition (compare SFr in Fig. 6C with MFR in Fig. 6B; P < 0.05). On the other hand, when movement was added to rhythmic low-force generation (MFr condition), flexion-phase work output was reduced by an amount no less than that in the rhythmic high-force, static or moving, conditions (compare MFr in Fig. 6C with SFR or MFR in Fig. 6B, both P > 0.05).
Compared with the nominal condition when the nonpedaling leg was stationary and relaxed (SR condition), no significant change in coordination of the pedaling leg during its flexion phase was detected when the nonpedaling leg generated a high or low tonic extensor force (SFT and SFt conditions). Neither was flexion-phase work output significantly changed (compare SFT and SFt in Fig. 6D with SR in Fig. 6A; both P > 0.05) nor muscle activity levels (Fig. 8, A-D, compare SFT and SFt with SR; all P > 0.05).
Although passive movement of the nonpedaling leg (i.e., MR condition)
did not change work output or muscle activity in the pedaling leg
during the flexion phase (quadrant 4, see preceding text), BF activity
was reduced during the limb extension-to-flexion transition phase
[quadrant 3; e.g., compare Fig. 7B (quadrant 3) with Fig.
7A (quadrant 3)]. BF iEMG activity over this transition region when the nonpedaling leg was moved was reduced by 24% compared with the nominal condition when the nonpedaling leg was stationary and
resting (P < 0.01). The addition of rhythmic
high-force generation by the nonpedaling leg (i.e.,
MFR condition) had no additional effect on BF
activity in quadrant 3 (P > 0.05). It should be noted that the limb extension-to-flexion transition (quadrant 3) is a region
of the crank cycle where BF is excited in forward two-legged pedaling
(Raasch et al. 1997; Ryan and Gregor
1992
).
Nonpedaling leg performance
The desired sensorimotor state of the nonpedaling leg was achieved
in most conditions as demonstrated by the force and EMG measures. When
the nonpedaling leg generated rhythmic force, the force trajectory
varied such that there was a single maximum and minimum per pedaling
cycle. The maximum occurred during the flexion phase of the pedaling
leg while the minimum occurred during the opposite phase (compare
phasing of pedaling leg force in Fig. 9A with phasing of maximum and
minimum nonpedaling force peaks in Fig. 9, B and
C). The level of extensor force attained and iEMG measured
increased with the level of the targeted force. Peak extensor force
levels were lowest during the no-force conditions (SR, MR),
intermediate during the rhythmic low-force conditions (SFr, MFr), and highest
during the rhythmic high-force conditions (SFR,
MFR; compare ,
, and
in Fig.
9B for maximum extensor force under static conditions, Fig.
9C under movement conditions). Correspondingly, iEMG levels
over the crank cycle were higher during high rhythmic force generation
(SFR, MFR) than during
no-force generation (SR, MR; Table 1).
Specifically, VM increased almost fourfold, which is significant
because it contributes the most to the force and power in limb
extension during pedaling (Raasch et al. 1997
). In
addition to the reduced force observed in the low-force condition, the
desired condition of reduced extensor performance was confirmed by iEMG
as VM decreased during low- (SFr,
MFr) compared with high-force generation
(SFR, MFR; Table 2).
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|
|
The only desired sensorimotor state not achieved well by the subjects
was tonic force generation in their nonpedaling leg (SFt, SFT). Instead of
being constant, the extensor force developed in the nonpedaling leg was
modulated over the cycle rather than constant (compare and
with
and
in Fig. 9D). The maximum force produced during
the tonic high-force condition (SFT) was within
the bounds of the "high target" region. If a true tonic force had
been maintained, the force minima (
and
, Fig. 9D) would have been equal to the force maxima (
and
, Fig.
9D). Although the force minima were significantly lower
(P < 0.01), the minimum force in the tonic conditions
was greater than the minimum force during the rhythmic conditions
(compare
and
in Fig. 9, D with B and
C, P < 0.01). Thus though substantial extensor force was generated by the nonpedaling leg throughout the
cycle in the tonic force conditions (SFt,
SFT), this force was phasically modulated, with
less force generated when the contralateral pedaling leg was in the
extension phase than the flexion phase.
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DISCUSSION |
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Equivalency of task mechanics
Since our experimental conditions were designed to maintain
consistent mechanical conditions in the pedaling leg, changes in
coordination of the pedaling leg are probably not due to differences in
task mechanics or pelvis motion. Pedaling-leg mechanics were highly
consistent among cycles because the torque generated by the constant
force spring was very consistent (Fig. 1A). Thus changes in
the pedaling performance of subjects could not be due to changes in the
mechanical load at the crank (cf. Ting et al. 1998b).
Although very little power is normally transmitted through the pelvis
in pedaling (Ingen Schenau et al. 1992
; Neptune
and Hull 1995
), energy transfer due to pelvis motion was
further minimized by a restraining hip belt. Even though pelvis motion
was slightly greater in movement conditions, this did not appear to
affect the decrease in flexion-phase crank torque as pelvis motion was not consistently found in either static or movement conditions. Thus
the decrease in flexion-phase crank torque was not an artifact of
pelvic motion.
Flexion phase pedaling coordination is modulated by contralateral-leg rhythmic extensor force generation
In a previous study, increased flexion-phase muscle
activity (and reduced negative work) compared with bilateral
pedaling was measured in subjects pedaling unilaterally with the
contralateral leg stationary and relaxed, even though the task
mechanics of the pedaling legs were identical in the unilateral and
bilateral pedaling conditions (Ting et al. 1998b).
Moreover, subjects did not compensate for this increase in muscle
activity when given visual feedback of crank torque profile. In our
current study, subjects pedaling unilaterally who generated a rhythmic
extensor force with their contralateral nonpedaling leg
reduced the amount of muscle activity (and
increased the amount of negative work) in the pedaling leg
during its flexion phase. The level of reduction of flexion-phase
muscle activity also depended on whether contralateral leg movement
accompanied the contralateral force generation. Because of major
technical limitations in the experimental apparatus at the time, it was
impossible for us to include a bilateral pedaling condition; thus we
are unable to directly compare the current unilateral pedaling
condition (the SR condition) with bilateral pedaling. Nevertheless, the
profile of crank torque generation and EMG patterns in the pedaling leg
in our current study resemble those of the unilateral pedaling
condition in the previous study (Ting et al. 1998b
).
Thus the mechanisms used by subjects in our current study to reduce
flexion-phase muscle activity during contralateral rhythmic extensor
force generation may indeed be those involved in the reduction of
flexion-phase muscle activity in bilateral pedaling. Regardless, the
two studies demonstrate the importance of contralateral sensorimotor
state in shaping the default locomotor pattern observed in the flexion
phase of pedaling.
Although our current study concentrated on the concurrent effects of
contralateral extension on the ipsilateral flexion phase, contralateral
force generation appears to affect other phases as well. For example,
when the contralateral nonpedaling leg was moved (MR and
MFR conditions), BF activity in the pedaling leg was reduced not only in its flexion phase but also in its
extension-to-flexion transition phase. In contrast, when the
contralateral leg was stationary and at rest (SR condition), BF
activity in the pedaling leg appears to be high not only in its flexion
phase (Fig. 8A) but in its extension phase as well. However,
the changes that may occur in phases other than flexion (i.e., the
extension-to-flexion and flexion-to-extension transition phases, and
the extension phase) (Raasch and Zajac 1999;
Raasch et al. 1997
) cannot be reliably interpreted for
the contralateral force generating conditions (cf.
contralateral-movement-only condition) because the experimental design
did not control the sensorimotor state of the nonpedaling leg well in
the other phases. Therefore we have focused on the effects found in the
flexion phase of the ipsilateral pedaling leg.
Phase-dependent responses to contralateral sensorimotor state
The different sensitivity of pedaling-leg flexion-phase BF and SM
activity and of RF and TA activity to contralateral extensor force
generation and movement may reflect their different biomechanical contributions to the execution of pedaling. Computer simulations and
experiments of forward and backward pedaling show that BF and SM act
primarily to accelerate the limb posteriorly and RF and TA the limb
anteriorly, regardless of pedaling direction (Neptune et al.
2000; Raasch and Zajac 1999
; Ting et al.
1999
). However, all four muscles (BF, SM, RF, TA) can
contribute secondarily to limb flexion (Neptune et al.
2000
; Raasch and Zajac 1999
; Raasch et
al. 1997
; Ting et al. 1999
), though in different
regions of the flexion phase. In forward pedaling, BF and SM contribute
to early limb flexion, near the end of their primary contribution to
posterior limb movement (e.g., the latter portion of quadrant 3 and the
initial portion of quadrant 4; Fig. 3). RF and TA contribute to late
limb flexion, near the initial portion of their primary contribution to
anterior limb movement (e.g., the latter portion of quadrant 4 and the
initial portion of quadrant 1; Fig. 3). Thus the effects of
contralateral sensorimotor state on muscles contributing to flexion
could reflect a general control strategy for flexors. On the other
hand, the effects may depend specifically on each muscle's primary
function. Since we were unable to record from iliacus, psoas, and the
short head of the biceps femoris, we are unable to assess the
contributions from those uniarticular muscles that primarily contribute
to limb flexion. Therefore though we hypothesize that uniarticular knee
and hip flexor muscles would also show strong influences from
contralateral movement and rhythmic force generation, we cannot rule
out an alternative hypothesis that only multifunctional muscles
contributing secondarily to flexion are affected.
The invariance in level of reduction of flexion-phase BF and SM
activity to contralateral sensorimotor state (Fig. 8, A and B: SFR, MFR,
SFr, and MFr conditions) is
consistent with a feedforward mechanism based on initiation of force
generation in the contralateral leg. Specifically, neither force level
nor movement modulated flexion-phase BF and SM activity. Instead,
contralateral force generation (both low and high) caused a uniform
decrease in flexion phase BF and SM activity in all subjects,
regardless of movement condition (Fig. 8, A and
B). Since BF and SM activity returns to baseline (Fig. 7)
before peak extensor force is generated in the contralateral leg (i.e.,
before ~315° of the pedaling-leg crank angle, Fig. 9, A
and B), the magnitude of flexion-phase BF and SM activity
may be triggered or programmed by feedforward signals related to
initiation of contralateral extensor force rather than modulated by
feedback related to contralateral force or movement amplitude.
Flexion-phase BF and SM activity may be subject primarily to
contralateral feedforward influence because their contribution to
propulsion is limited to early flexion phase (Raasch et al.
1997). Similarly, when the leg is obstructed during early swing
phase of gait, an invariant muscle coordination pattern emerges; yet
when the obstruction occurs in late swing, the pattern is more variable
(Eng et al. 1994
, 1997
). Further, in human locomotion, it appears that flexor activity may be triggered by peripheral input
with activity levels modulated by central mechanisms (Dietz 1992
). However, there could have been other influences acting on BF and SM activity that were unmeasurable due to the overall low
level of EMG activity in BF and SM.
In contrast to BF and SM, flexion-phase RF and TA activity may be
influenced by contralateral feedback related to successful generation
of both force and movement because the reduction in RF and TA activity
only occurred when contralateral force generation accompanied
contralateral leg movement, which is the condition (MFR) studied by us that most replicates
two-legged pedaling. Further, high inter- and intra-subject variability
of RF and TA activity was found in the force-generating conditions. For
example, an increase of RF and TA activity, the opposite
effect from the norm, was measured in some subjects during isometric
force generation (compare LRF and LTA in Fig. 7C to LRF and
LTA in Fig. 7A). Such an increase is consistent with the
necessity to excite both contralateral extensors and ipsilateral
flexors if contralateral extension alone inadequately propels the
crank, such as against very large loads (Raasch et al.
1997). Thus RF and TA activity may be modulated both positively
or negatively by contralateral feedback. However, the precise state
associated with the successful generation of a high-force and
anti-phase movement during contralateral extension appears to cause a
decrease in RF and TA activity. Because a precise correlation of RF and
TA with force level was not found, RF and TA activity may be affected
by other, uncontrolled variables in the experiment. Further, the
relatively large range of permissible "high" forces may have
contributed to the lack of statistical sensitivity in some conditions.
Neural interlimb coupling constraints on muscle coordination as
evidenced by phase-dependent responses to contralateral state appear to
reflect functional similarities between ipsilateral flexors and
contralateral extensors in two-legged pedaling. In moderate load and
steady-state conditions, flexion-phase muscle activity is relatively
low while contralateral extensors provide the majority of propulsive
power (Kautz and Hull 1993). Although the maximal
contribution to propulsion by flexors is small compared with the
contribution of extensors, flexors must be recruited nevertheless under
high load conditions (Raasch et al. 1997
) or when
extensors generate inadequate force (e.g., hemiparetic subjects in
Kautz and Brown 1998
). Thus it is not surprising that a
mechanism for activating flexors might exist that is based on
contralateral extension that parallels the potential biomechanical
contribution of flexors to pedaling. The proposed feedforward and
feedback mechanisms would cause flexor activity to increase if either
contralateral extension is not initiated or contralateral extension
produces inadequate force or movement. Because extrinsic loading in the pedaling leg was constant, contralateral phasic motor commands and
dynamic force and movement feedback signals are possible candidates involved in neural interleg coupling. The neural interleg coupling pathways identified here to be operational during the flexion phase may
be indicative of a bilateral recruitment strategy for crank propulsion.
However, since activity in uniarticular knee and hip flexors was not
measured, the effect may be characteristic of only muscles that act
primarily during the limb extension-to-flexion transitions, such as
ankle muscles (TA) and biarticular thigh muscles (RF, BF, and SM), that
have a secondary contribution to flexion (cf. Cabelguen et al.
1981
; Perret and Cabelguen 1980
).
One legged pedaling may excite contralateral pattern-generating elements
While the conditions of the nonpedaling leg affect pedaling
coordination, conversely, sensorimotor control of one-legged pedaling may provide excitatory drive and modulate pattern-generating elements of the contralateral (nonpedaling) leg. Although the neuronal basis of
pedaling coordination is unknown, phasic reflex modulation patterns
during pedaling are similar to reflex patterns during walking; thus
common neural elements may be employed in pedaling and walking
(Brooke et al. 1992-1995
, 1997
; Collins et al.
1993
; McIlroy et al. 1992
) In our current study,
rhythmic muscle activity patterned as in two-legged pedaling was evoked
in the passively moved nonpedaling leg even though subjects were
instructed to relax the leg (Fig.
10B, MR condition). The
phasing of the low-amplitude muscle activity in the "relaxed"
nonpedaling leg was similar to the phasing of activity in a leg that
actually pedals (compare Fig. 10B with Fig. 7A).
The expression of rhythmic, pedaling-like activity in a nonpedaling leg
may depend on the presence of pedaling in the other leg because in the
control condition (MRØ), where the ipsilateral (usually pedaling) leg
was stationary and relaxed, very little pedaling-like EMG activity was
evoked in the contralateral servomoved leg (Fig. 10A). The
expression of rhythmic, pedaling-like activity in a nonpedaling leg may
also require movement because no pedaling-like pattern was evoked in
the static, relaxed nonpedaling leg when the other leg pedaled (SR
condition) (unpublished nonpedaling-leg EMGs). On the other hand, it is
possible that unilateral pedaling without contralateral movement does
indeed excite contralateral pattern generating elements but the
expression of the pattern in the motor output is below threshold.
Further, the pattern of muscle activity in the nonpedaling leg during
the rhythmic extensor force condition (MFR) was
consistent with the extension-phase two-legged pedaling pattern. In
contrast, in the absence of pedaling when subjects practiced generating
rhythmic force, a nonpedaling-like muscle activity pattern was
observed. All of these observations are consistent with bilateral
pattern generating elements being excited and modulated even when only
one leg pedals, though expression of the pattern in the nonpedaling leg
may depend on its sensorimotor state.
|
The fact that subjects could not produce a constant force with the
nonpedaling leg while the other leg pedaled further supports the notion
that sensorimotor control of one-legged pedaling excites the pedaling
pattern-generating elements of the contralateral leg. Rather than
produce a constant force throughout the crank cycle, subjects produced
less force in the nonpedaling leg as the pedaling leg executed its
extension phase. Generation of the phasic motor output in the
nonpedaling leg (evidenced by EMGs as well as force) could be caused by
tonic afferent feedback acting to augment the centrally generated
locomotor rhythm (review, Rossignol 1996). Phasic
motor-output generation in the nonpedaling leg could also be the result
of rhythmic inhibition of motoneurons due to activation of spinal
locomotor circuits by the other pedaling leg (Orsal et al.
1986
) in parallel with tonic supraspinal excitation of the
motor pools in the generation of the desired tonic force output.
Alternatively, the rhythmic afferent signals from the pedaling leg
could have activated interlimb spinal pathways associated with
reciprocal inhibition pathways that interact with the tonic supraspinal
command at the spinal level (for review, see Jankowska and
Edgley 1993
). Finally, long-loop sensorimotor pathways could have phasically modulated the descending command (Asanuma and Keller 1991
).
Bilateral sensorimotor signals modulate the locomotor pattern
This study provides evidence of modulation of flexor activity in
bifunctional muscles by contralateral sensorimotor signals during a
steady-state pedaling task. Complex interlimb influences have also been
noted in human walking during obstacle avoidance and tripping where
only one leg is perturbed (Dietz et al. 1986; Eng
et al. 1994
). Although neuronal interlimb coordination
mechanisms no doubt were present in these studies, the effects due to
neural interlimb coupling could not be isolated because of
instantaneous mechanical transmission of force to the nonperturbed leg
that accompanies contralateral limb acceleration (Yamaguchi and
Zajac 1990
; Zajac 1993
). In our study,
mechanical interlimb coupling was almost completely eliminated. Because
movement or force generation in one limb did not mechanically affect
the other limb, the effect of extension-phase movement and force
generation in one leg on flexion-phase EMG patterns in the other leg
must be due to neuronal interlimb coupling.
While the basic rhythm and pattern of muscle activity during locomotion
may indeed be generated through traditional pattern generation
elements, sensory inflow, including motion-dependent, and
task-dependent feedback, also affect the relative timing and duration
of muscle activity, as shown in cats (Cabelguen et al. 1981; Smith et al. 1993
), chicks (Bekoff
et al. 1987
), and humans (Andersson et al.
1997
). A traditional excitatory connection between contralateral extensor and flexors cannot by itself reproduce the key
features of the basic locomotor pattern of flexor bursts being shorter
in duration than extensor bursts (cat, Rossignol 1996
;
chicks, Bekoff et al. 1987
) and of the flexor burst
period and swing duration changing little with speed while extensor
bursts change greatly (cat, Cabelguen et al. 1981
;
Smith et al. 1993
; chick, Bekoff et al.
1987
; humans, Andersson et al. 1997
;
Nilsson et al. 1985
). Even in fictive cat preparations,
sensory inflow can significantly change the locomotor pattern
(Cabelguen et al. 1981
; Perret and Cabelguen
1980
). In spinal cats, unilateral deafferentation disrupts
bilateral pattern generation (Giuliani and Smith 1987
). In chicks, bilateral deafferentation of the legs causes flexor activity
to increase to nearly the same length as extensor activity (Bekoff et al. 1987
), further supporting the concept
that sensory information preferentially shapes flexion-phase activity.
Finally, the variations in the locomotor pattern observed during
different task conditions such as speed and incline (Andersson
et al. 1997
; Carlson-Kuhta et al. 1998
;
Nilsson et al. 1985
; Pierotti et al. 1989
; Smith et al. 1993
, 1998
) point to the
final motor output being shaped by task- and motion-dependent signals
to adapt to task demands.
Thus a rich repertoire of task-dependent neuronal interlimb coupling
mechanisms, which tend to coordinate the legs as a functional unit, may
modulate the basic locomotor pattern during steady-state locomotion and
in response to small perturbations. Further, feedforward and feedback
mechanisms appear to modulate EMG patterns in pedaling (McIlroy
and Brooke 1987) as well as walking (Dietz et al.
1986
; Eng et al. 1994
), with some EMG patterns
changing stereotypically and others in proportion to the stimulus. In
addition, these EMG responses vary considerably with task mechanics as
well as the phase of the stimulus. For example, the EMG response
elicited when the swing leg is impeded by its striking an obstacle
(Eng et al. 1994
, 1997
) differs greatly from the
response when the swing leg is impeded by a rope instead (Dietz
et al. 1986
). In our current study, inadequate force generation
in extension by one leg appears to increase flexor force generation in
the other leg. Similarly, in hemiplegic chicks, the extension phase of
the paretic limb is accompanied by an increased flexion in the
contralateral leg, causing an asymmetrical gait pattern (Muir
and Steeves 1995
; Muir et al. 1998
). Thus many
interlimb coordination patterns probably exist to coordinate the legs
as a functional unit, with the specific set of muscles affected and the
intensity of each effect depending on the exact bilateral afferent and
efferent state.
Such interlimb coupling mechanisms, which serve to modulate muscle
activity during steady-state conditions, may be distinct from the large
repertoire of task- and state-dependent interlimb coordination reflexes
that reset or interrupt the locomotor rhythm. Studies have demonstrated
that stimulation of cutaneous and proprioceptive afferents in various
cat preparations prolong and enhance the extension phase and the
contralateral flexion phase in a manner similar to that of a
Sherringtonian flexion reflex (Duysens and Pearson 1976,
1980
; Guertin et al. 1995
; Pearson et al.
1992
). Further, improper ground support (Gorassini et
al. 1994
; Hiebert et al. 1994
) or stimulation of
flexor reflex afferents (Schomburg et al. 1998
)
in spinal and intact cats during extension causes a rapid ipsilateral
flexion and a contralateral extension. Although such responses do serve
to coordinate the legs in a functional manner against large
perturbations, they tend to truncate or reset the locomotor rhythm and
could be considered separate from the class of interlimb coordination
mechanisms demonstrated in our study.
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ACKNOWLEDGMENTS |
---|
We thank M. Van der Loos for design and implementation of the motor apparatus and data collection system, C. Dairaghi for data collection expertise, and J. Holt for help both collecting and analyzing data.
This research was supported by National Institute of Neurological Disorders and Stroke Grant NS-17662 and the Rehabilitation Research and Development Service of the Dept. of Veterans Affairs (VA).
Present addresses: L. H. Ting, Neurological Sciences Inst., 1120 NW 20th Ave., Portland, OR 97209; D. A. Brown, Programs in Physical Therapy, Northwestern University Medical School, 645 N. Michigan Ave., Suite 1100, Chicago, IL 60611-2814.
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FOOTNOTES |
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Address for reprint requests: F. E. Zajac, Rehabilitation R and D Center (153), VA Palo Alto Health Care System, 3801 Miranda Ave., Palo Alto, CA 94304-1200.
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 30 June 1999; accepted in final form 24 February 2000.
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REFERENCES |
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