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JOURNAL OF APPLIED BIOMECHANICS, 1997,13,334-346
O 1997 by Human Kinetics Publishers, Inc.
Changes in Resultant Pedal
Reaction Forces Due to Ankle
Immobilization During Pedaling
Cheryl D. Pierson-Carey, David A. Brown,
and Christine A. Dairaghi
The purpose of this study was to determine the effects of limiting ankle motion on
pedal forces. Sixteen adults pedaled an instrumented ergometer against constant cadence and frictional load while wearing hinged braces. Ankle motion was limited under four randomly assigned conditions: both braces unlocked (UL), only the preferred
leg (PL) brace locked, only the nonpreferred leg (NPL) brace locked, and braces on
both legs (BL) locked. Measurements of pedal force, crank, and pedal angles were
sampled at 2001s for 20 s. With both braces locked, resultant force mean magnitude
decreased during the downstroke, due to reduced radial crank force. Asymmetry between PL and NPL decreased during the power phase when only PL was braced but
increased when only NPL was braced. It was concluded that constrained ankle motion, as may occur with ankle injury or hemiplegia, reduces the ability to transmit
power during the downstroke while enhancing ability during the upstroke.
Propulsion during human locomotion requires a firm interface between the foot and
the environment. Effective propulsion relies largely upon the ability of the ankle joint to
place the foot in a stable position (Rodgers, 1988). Consequently, biomechanical and/or
neurological impairments at the ankle may result in ineffective propulsive capabilities.
When these impairments are unilateral, an asymmetry in propulsive forces may lead to
unstable and unsafe locomotion (Giuliani, 1990). Ankle immobilization or disuse of the
ankle joint secondary to paresis and/or muscle weakness can lead to viscoelastic and noncontractile tissue changes (Cornwall, 1984; Dietz & Berger, 1983; Giuliani, 1990; Smidt,
1990). This biomechanical impairment is manifest as a reduction in or complete lack of
mobility at the ankle, which may place the foot in positions that are less effective during
the propulsive phases of locomotion.
To study the relative contributions of ankle joint mobility to locomotor propulsion,
we used a task that has well-defined and well-controlled mechanical constraints. Walking is a difficult locomotor task to study, because people can demonstrate large
interindividual variations, and patients may have protective gait mechanisms that are hard
Cheryl D. Pierson-Carey was a graduate student at Samuel Merritt College, Oakland, CA,
when this study was conducted. She is currently with the Department of Physical Therapy at Kaiser
Permanente, Fremont, CA 94538. David A. Brown and Christine A. Dairaghi are with the Rehabilitation Research and Development Center, Veterans Affairs Health Care System, 3801 Miranda Ave.,
Palo Alto, CA 94304. Direct correspondence to David A. Brown.
Pedal Reaction Forces
335
to experimentally control (Benecke, Conrad, Meinck, & Hohne, 1983). The variables related to bipedal movement during pedaling, however, are easier to control, since the kinematics of the legs are constrained by the bicycle's crank trajectory. A bicycle ergometer, modeled as a planar five-bar linkage system with only three degrees of freedom, is
mechanically simple and lends itself to quantitative analysis (Gonzalez & Hull, 1989;
Hull & Jorge, 1985).
Pedaling studies provide valid and reliable measurements that support the use of
bicycle ergometer pedaling as an ideal way to study the biomechanical components of
leg movement (Fregley & Zajac, 1996; Gonzalez & Hull, 1989; Hull & Jorge, 1985).
Environmental conditions, such as seat height or footlpedal interface, and task conditions, like workload or velocity, can be manipulated for systematic study of functional
locomotor forces, kinetics, and kinematics (Gregor, Broker, & Ryan, 1991; Kautz, Feltner,
Coyle, & Baylor, 1991). Bicycle ergometer pedaling has been used to study asymmetrical force production in both unimpaired populations (Daly & Cavanagh, 1976; Ericson,
Ekholm, Svensson, & Nisell, 1985; Ericson, Nisell, & Nemeth, 1988) and patient populations (Benecke et al., 1983; Brown, Burgar, Kautz, Dairaghi, & Dunn-Gabrielli, 1994;
Brown, Dairaghi, Stevenson, Wu, & Zajac, 1992; Giuliani, Harro, & Rosecrance, 1989;
Rosecrance & Giuliani, 1991). Furthermore, bicycle ergometers are readily available in
clinics and are commonly used for strengthening and endurance exercise in many types
of rehabilitation.
The foottpedal interface, by way of the crank, is the firm link with the environment
for propulsion during pedaling. Ankle position may significantly affect the application of
forces to the crank produced by the muscles of the leg. Likewise, the ankle trajectory is
known to change as a result of changes in workload (Ericson et al., 1988). "Ankling," a
technique used by world-class cyclists that is thought to increase effective forces to the
crank, can increase crank acceleration (Gregor et al., 1991). Studies with neurologically
unimpaired subjects show somewhat asymmetrical kinematics of bipedal tasks that might
be related to the asymmetry of forces produced by each leg in the absence of ankle impairments (Benecke et al., 1983; Daly & Cavanagh, 1976).
Abnormal ankle trajectories may alter the biomechanics of pedaling and impair
performance. Studies of subjects with neurological impairments suggest that decreased
ankle range of motion may produce asymmetrical forces during bicycle ergometer pedaling (Benecke et al., 1983; Rosecrance & Giuliani, 1991). In a study by Benecke et al.
(1983), the torque curve profiles produced by 12 subjects with marked spastic paraparesis
and deficits in ankle dorsiflexion showed a maximum peak torque occurring at a different
pedal angle compared to the other subjects in the study. The authors concluded that the
ankle dorsiflexion joint deficits created an unfavorable angle of application of forces to
the pedal. Rosecrance and Giuliani (1991) identified abnormal ankle kinematic trajectories during the upstroke phase of the pedaling cycle in a group of subjects with hemiplegia. In contrast, their neurologically unimpaired subjects produced a predictable dorsiflexion~~lantar
flexion ankle pattern during the same cycling phase. Although these two
studies singled out ankle joint mobility as a probable important contributor to the altered
pedaling forces, the direct relationship between ankle movement and pedal forces was not
assessed.
In the present study, we used neurologically unimpaired subjects to determine the
effects of severely limited ankle motion on pedal reaction forces during steady-state ergometer pedaling. Since limited ankle mobility is thought to be a major contributor to impaired pedaling performance, we hypothesized that bracing the ankles in a neutral position would significantlychange the pedal reaction force trajectories. Further, we expected
Pierson-Carey, Brown, and Dairaghi
336
that bracing only the ankle of the leg that produces the majority of force would reduce the
overall asymmetry of forces produced by the two legs, while increasing asymmetry of
forces by limiting ankle motion of the opposite leg.
Methods
Subjects
+
Sixteen subjects (8 female, 8 male) between the ages of 18 and 35 years (mean = 28 3.5
years) were selected for this study. In an effort to eliminate force asymmetries that could
be due to anthropometric andlor passive tissue differences, subjects were initially screened
for abnormal ankle range of motion in either of the two legs (Daniels & Worthingham,
1986) and leg length discrepancies of greater than one-half inch (Subotnick, 1976). Subjects were also screened for signs, symptoms, or significant history of neurological or
cardiovasculardisease. To avoid including subjects who had learned to adopt skilled ankling
strategies (Gregor et al., 1991), we chose only recreational cyclists (operationally defined
as anyone who bicycled two or fewer times per week).
The protocol for this study was approved by the Stanford University Human Subjects committee and was in compliance with federal policy on the protection of human
subjects. Informed consent was obtained from each subject before the experiment began.
A bicycle with an instrumented pedal system, a seat with a backrest, and a workload
ergometer was used (Figure 1). Pedal force transducers, composed of strain gauges in a
1 1
AID
Unlocked Anre Brace
Pedal Forces
Pedal Encoder Signal
-
Figure 1
Experimental setup showing the bicycle ergometer, subject, and measurements
recorded. Shown is a schematic representation of the unlocked (right) and locked (left) brace
conditions.
Pedal Reaction Forces
337
Wheatstone bridge, measured the normal and fore-aft forces on each pedal. Forces were
measured with an absolute error bound of f 5 N (Newmiller, Hull, & Jajac, 1988). Three
optical encoders, one at the end of each pedal spindle and one coupled to the crank, provided measurements of the relative crank and pedal angles throughout 360" of rotation.
The seat and seat post were attached to a backrest, oriented 20" from the vertical, that had
a semirigid shoulder harness and belt system. Handlebars were removed from this bicycle
because subjects were able to lean safely and comfortably against the backrest. The absence of handlebars ensured that measured pedal forces were not influenced by the stabilizing forces generated from the upper body and arms.
The workload ergometer was adapted from a standard Monark ergometer. A freewheel allowed the flywheel to spin if the subject slowed down or ceased pedaling. The
friction on the flywheel, indirectly measured by a load cell, was controlled by adjusting
the belt tension. To provide an output measure of velocity, an in-house-built electronic
circuit was used to differentiate the crank position encoder signal. This signal was visually
displayed to the subjects during pedaling.
Two Ultra 3 ROM ankle foot orthoses (Orthopedic Technology, Inc., Tracy, California), referred to as "braces" in this study, were modified for use and worn by each
subject throughout the testing. Each brace had an adjustable locking hinge at the malleoli
with rigid medial and lateral uprights that spanned the length of each shank. Additionally,
each brace had an inner toe-to-knee liner, an open-toe cushioned rubber sole, and six
Velcro closure straps. The braces were modified by attaching cycling cleats to the bottom
just beneath the forefoot so that the foot could be positioned over the pedal in the conventional way. These cleats clipped into a locking system on each pedal that ensured a rigid
footlpedal connection (Davis & Hull, 1981).
Experimental Protocol/Procedures
Subjects pedaled the ergometer under four different braced conditions. First, subjects pedaled while both braces were unlocked (UL) so we could assess the control condition profile of forces and their relative asymmetry when the ankles were allowed to move freely.
The other three conditions involved locking the brace of either the preferred leg (PL), the
nonpreferred leg (NPL), or both legs (BL) at a 90" angle (neutral position). This locking
allowed us to determine the kinetic effects of ankle immobilization during asymmetrical
(PL or NPL) or symmetrical (BL) kinematic conditions. The order of the three locked
brace conditions was randomly assigned for each subject. The neutral position was chosen
because it placed the ankle muscles at functional lengths (i.e., neither too long nor too
short) that avoided possible strain and damage to the muscles, ligaments, and joints. All
testing was done on the same day for each subject.
Our operational definition of range of motion was the anatomical movement available at a joint from a defined zero starting position as measured by a goniometer. For the
purposes of this study, we were interested in ankle dorsiflexion (DF) and plantar flexion
(PF) motion, which takes place between the talus and tibia and fibula within the ankle
mortise. The neutral position occurs when the ankle is at a right angle to the lower leg and
the lower leg is at a right angle to the thigh; this is also the zero starting position for
measurements of DF and PF. The rigid uprights of the ankle brace were aligned with the
shaft of the fibula so that when the brace was locked at 90°, the ankle was placed in a
neutral position. Goniometric and visual inspection was used to ensure proper position of
the ankle in the brace.
Each trial consisted of the subject pedaling the ergometer against a 15 N . m frictional load, at a rate of 60 rpm, for approximately 30 s. Data collection began after the
Pierson-Carey, Brown, and Dairaghi
338
subject reached the target speed and maintained that speed with some consistency (approximately 10 s). Measurements of pedal force, crank, and pedal angles were collected
during each cycle. Measurements were analog-to-digital ( A D )converted and sampled at
2001s for 20 s. The digitized samples were stored on a 386 personal computer for further
reduction and analysis.
Data Analysis
Tangential (Ft) and radial (Fr) component crank forces were calculated using the pedal
reaction forces and pedal and crank optical encoder measurements. The Ft component of
crank force is defined in this paper as that component of force directed perpendicular to
the crank arm,whereas the F r component is defined as that component directed parallel to
(i.e., along the length of) the crank arm. Resultant crank forces (Rfl were calculated as the
vector sum of the Ft and F r components at each crank position. For purposes of characterizing force asymmetry, we defined the preferred leg (PL) as the leg that produced more
than 50% of the mean Ft component during the UL condition when both ankles moved
freely. The other leg was referred to as the nonpreferred leg (NPL).
Data were analyzed by first inspecting the individual pedaling trials of all subjects.
Then, cycles were averaged for each pedaling trial across subjects. After the average work
done by both legs for each pedaling condition was calculated, all force measures were
divided by this value. This process allowed all force measures to be expressed as a percentage of the average overall (mean) work done; hence, we could account for any differences in workload due to uncontrolled tightening or loosening of the friction belt. The
average total work done by both legs was computed for each pedaling trial and then used
to normalize the Ft and F r components and the Rf values so that they could be grouped
with other subjects by condition (e.g. UL, PL, NPL, and BL). Finally, the force data were
normalized to the mean Ft component value for that trial and grouped with other subjects
by condition. In order to compare the magnitude and direction of forces, data were averaged over a 30' crank angle range, and the resulting average value was compared. When
the relative asymmetry of forces was compared between the PL and NPL, data were first
averaged over the 30" window, and then the differences between the two legs were calculated.
Repeated-measures one-way ANOVA, blocked by subject, was used to test the hypothesis that differences in Rf, Ft, and Fr component crank forces existed across the locked
brace conditions (Kvanli, 1988). Significance was set a t p < .05 for all tests, and a Tukey
post hoc analysis was used to test for multiple comparisons among the conditions (Kvanli,
1988).
Results
Magnitude and Orientation of Reaction Forces
The time history of magnitude and orientation of the Rf were altered as a result of locking
the braces of both ankles in a neutral position (BL condition). The Rf magnitude and
direction at each of 12 regions of the crank cycle are displayed in Figure 2. Throughout
most of the crank cycle, with the exception of the top-dead-center (0' or 360') region, the
Rf magnitude decreased as a result of locking the braces. The biggest difference in magnitude occurred at the bottom-dead-center (180') region. The orientation did not appear to
change significantly, except during the 270' and 300" regions, and this was only by small
amounts (5' and 8' difference, respectively). Results were similar for the PL and NPL
Pedal Reaction Forces
0 deg
TDC
270 deg
BDC
180 deg
KEY: 1 unit = mean F t component value
-
-
p<.05 Rf magnitude difference
b = p<.05 Rf orientation difference
a
unlocked
both locked
Figure 2 - Resultant crank force (Rf)vector magnitude and orientation at each of 12 regions
of the crank cycle comparing the UL (both braces unlocked) condition with the BL (both braces
locked) condition. Rf was averaged across all subjects, and data are from the preferred leg
(PL) only. (Results from the nonpreferred leg [NPL] were similar to those from the PL). Rf was
significantly reduced @ < .05) over the 120" to 300"regions. TDC = top dead center relative to
the seat tube (0" region); BDC = bottom dead center relative to the seat tube (180" region).
conditions, such that reductions in magnitude and changes in orientation of Rf were altered in the locked brace ankle. Thus, limiting ankle mobility helped reduce the forces
applied to the pedal.
The Ft and Fr components of the Rf also showed statistically significant changes in
magnitude as a result of locking the braces. The Ft component, that is, the component
which is perpendicular to the crank and, thus, responsible for accelerating the crank in a
forward direction, decreased during the 120" and 150" regions of the downstroke phase,
yet increased (i.e., became less negative) during the 210" through 300" regions of the
upstroke (Figure 3). The net effect of these changes was a relatively constant mean Ft
component, or total work, done by each leg throughout each cycle and across all test
conditions. The percentage of mean work done by each leg was 53.9% (PL) versus 46.1%
k 3.0% (NPL) and remained unchanged across conditions O, = .7). Therefore, it appears
that the net effect of locking the braces was minimal, as reflected by the Ft component,
mainly because the loss in force transmittal during the downstroke was made up by an
increase in force transmittal during the upstroke.
The Fr component, that is, the component which is directed along the crank and,
therefore, does not directly contribute to crank acceleration, decreased from the 150" re-
Pierson-Carey, Brown, and Dairaghi
+ unlocked
- - o--
both locked
crank angle (degrees)
Figure 3 -Mean tangential crank force (Ft) component of the resultant force across all subjects
and at each of 12 regions of the crank cycle for the UL (both braces unlocked) and BL (both
braces locked) bracing conditions. Mean Ft was decreased with the BL condition during the
120' and 150°regions yet was increased (became less negative) during the 210° to 300° regions.
gion through the 300" region of the crank cycle (Figure 4). This decrease occurred during
the region of the crank cycle when the Rf was experiencing its greatest reduction during
the locked brace conditions. Therefore, since the Fr decrease was not made up by an
increase anywhere else in the crank cycle, it appears that the reduction in the Fr component was the major contributor to the reduction in the Rf trajectory.
Asymmetry of Reaction Forces
Since locking the braces of both ankles may effectively reduce the degrees of freedom of
the pedaling task, a reduced amount of relative force asymmetry between the two
legs under the BL condition might be expected. Indeed, the Ft component difference produced by the two legs was significantly reduced by 15% when both braces were locked
( p = .0036). This reduction in asymmetry was seen throughout most of the crank cycle
(Figure 5).
Subjects pedaled more symmetrically when only the PL brace was locked and less
symmetrically when only the NPL brace was locked (Figure 6). The difference between
the two conditions was most pronounced at about 150°, toward the end of the downstroke
phase. These results follow from the observation that the locked conditions reduced the Rf
applied to the crank (e.g., see Figure 2). Therefore, the asymmetrical characteristics of
these force profiles, caused by relative preference for one leg over the other, could be
altered by simply locking the brace of either the PL or the NPL.
Consistent with the results presented in the previous section on magnitude and orientation of reaction forces, percentage of mean work done by each leg remained unchanged under asymmetrical locking conditions when compared to the symmetrical UL
Pedal Reaction Forces
unlocked
-- -
both locked
0-
* p<0.05
crank angle (degrees)
Figure 4 -Mean radial crank force (Fr) component of the resultant force averaged across ail
subjects and at each of 12 regions of the crank cycle for the UL (both braces unlocked) and BL
(both braces locked) bracing conditions. Mean Fr was significantly different @ < .05) with the
BL condition during the O0 to 60' and 150' to 300° regions.
-- 0-
unlocked
both locked
crank angle (degrees)
Figure 5 -Mean differences in Rf values between the preferred leg (PL) and nonpreferred leg
(NPL) averaged across all subjects and at each of 12 regions of the crank cycle for the UL (both
braces unlocked) and BL (both braces locked) bracing conditions. Although there were no
significant changes found at each crank cycle region @ > .05), the average value for the BL
condition showed a 15% decrement in asymmetry across the entire cycle @ = .0036).
Pierson-Carey, Brown, and Dairaghi
--c--
preferred locked
both locked
---.o-.-.
nonpreferred locked
p<0.05
crank angle (degrees)
Figure 6 -Mean differences in Rf values between the preferred leg (PL) and nonpreferred leg
(NPL) averaged across all subjects and at each of 12 regions of the crank cycle for the BL (both
braces locked), PL, and NPL bracing conditions. There were significant differences < .05)
throughout the entire crank cycle, with the NPL condition demonstrating the least difference
and the PL condition demonstrating the greatest difference.
and BL conditions ( p = .7). This result further provided evidence for a conservation
of relative effort between the two legs that is independent of the biomechanical constraint.
Discussion
Limited Ankle Motion and Reaction Forces
One major result of this study is that the magnitude of reaction force trajectories was
reduced when ankle mobility was limited by locking the brace in a neutral position during
pedaling. Resultant forces were mostly reduced during the propulsive downstroke phase
of the crank cycle. Recall that resultant crank forces (Rfi are the sum of the Ft and Fr
trajectories at each position of the crank throughout the 360' cycle. During the initial
downstroke phase (0"-90°), the magnitudes of any changes in the Ft and Fr components
were very small and resulted in little change in the Rf trajectory. However, both the Ft and
Fr components were reduced during the terminal downstroke phase (90'-180°), accounting for the majority of Rf reduction observed during the downstroke. Since terminal
downstroke coincides with peak plantar flexion ankle angle (e.g., approximately 10" during unbraced ankle conditions; Brown, Kautz, & Dairaghi, 1996), the limited ankle plantar flexion during the locked brace condition might be responsible for the reduction in
forces. Therefore, the normal plantar flexion that occurs during unconstrained pedaling
might be a strategy for effectively transferring power from the hip and knee muscles to the
foottpedal interface.
Pedal Reaction Forces
343
Resultant forces continued to be reduced, as a result of locking the braces, during
the upstroke phases, although to a lesser extent. During the initial upstroke (180"-270°),
the majority of the reduction came from reduced Fr component contributions, with very
little change to the Ft component. However, during the terminal upstroke phase (27O0360°), most of the reduction in Rf, mainly from reductions to the Fr component, was
minimized by an increase in Ft component forces. This increase in Ft might have been due
to added stiffness at the ankle that was present when the brace was locked, thus allowing
leg muscles to be more effective in lifting the leg (and hence the pedal and crank) up
against gravity using the toeclips.
This added stiffness is usually supplied by ankle dorsiflexor muscle activation during unbraced pedaling and results in the ankle maintaining a relatively neutral position
despite the tendency for the ankle to plantar flex when pulling up against toeclips. Hence,
it appears that the net effect of locking the ankle brace in neutral is to limit the ability to
transmit Ft during the downstroke (reduce acceleration due to muscle activity and the
weight of the leg) by limiting the ankle's ability to place the foot and pedal in a more
effective plantar flexed position. Conversely, locking the ankle brace can enhance the
ability to transmit Ft during the upstroke (reduce deceleration due to the weight of the leg)
by enhancing stiffness at the ankle during this critical part of the crank cycle.
Asymmetry of Reaction Forces
The other major result of this study was that differences between reaction forces generated by each leg were reduced by locking the brace of the PL and increased by locking the
brace of the NPL. Recall that the PL produces the majority of Ft when both braces are
unlocked. Hence, when the PL brace was locked, the PL was less able to generate reaction forces throughout the cycle and the differences between it and the NPL were
reduced. Consequently, when the NPL brace was locked, the reaction force differences
increased. Finally, since the BL condition tended to constrain each ankle equally, and
since the reaction force differences between both legs were coincidentally reduced, there
may be a causal relationship between ankle motion asymmetry and pedal reaction force
asymmetry. Because we were unable to measure ankle motion directly during the unconstrained pedaling in this study, it remains for further studies to correlate ankle motion
asymmetry with reaction force asymmetry during unconstrained pedaling. However,
this is not to suggest that reaction force asymmetries are entirely due to mechanical constraints at the ankle.
Sensory feedback changes during the locked brace conditions might be partly responsible for triggering adaptations in the neural motor control strategy used to produce
consistent muscle forces and work output by each leg. Central nervous system (CNS)
integration in neurologically intact individuals enables them to use continuous sensory
feedback to make necessary and continuous adjustments in their motor output. Leg movements are a good case in point. Thelan (1992) showed that coordination between the two
legs is dynamic; that is, the movement status of each leg transmits information through
neural interconnections, thus determining synchronized leg movement. Her work with
newborns and young infants underwater versus in midair demonstrated this ability of the
CNS to both detect and respond to changes in the environment.
Another component of the neural control of locomotor behavior has been attributed to central pattern generators (CPGs) (Grillner, 1981). Central pattern generators initiate and control the spatial and temporal output of motoneurons during rhythmic, repetitive leg movement, such as pedaling. Although CPGs in the spinal cord allow the legs to
Pierson-Carey, Brown, and Dairaghi
344
act independently, a coupling between the CPGs results in coordinated, reciprocal
movement. Perhaps this CPG coupling enables a normal neurological system to find and
maintain a comfortable and relatively symmetrical movement pattern under normal pedaling conditions (free ankle motion) and adverse pedaling conditions (constrained ankle
motion).
The pedaling setup used in the present study differs, in several ways, from previous
pedaling studies that assessed asymmetry of pedal forces in neurologically unimpaired
populations. First, the backrestheat configuration inclined subjects backward, placing their
hips in a more extended position, thus altering the mechanics of pedaling. Second, the
absence of handlebars did not allow subjects to stabilize their upper bodies by leaning
forward. Thus, upper body forces, present in conventional cycling, were not included.
Finally, the lower leg braces, used in this study, added equal mass to the subjects' legs and
increased the height of the fooupedal interface.
In spite of the differences in the pedaling setup used, the pedal force asymmetry
found in this study appears quite similar to that found in previous studies using conventional pedaling setups with neurologically unimpaired subjects. Daly and Cavanagh (1976)
reported that peak crank torques averaged 54% versus 46% of the total peak torque for
each leg, which is in agreement with the results presented in this paper. In addition, typical
crank torque curves published elsewhere have similar features to those presented in this
study (see Gregor et al., 1991, for review). In summary, although the pedaling setup in this
study is different, the resulting kinetics of pedaling do not appear to be different from
those of more conventional pedaling studies.
Conclusions
The results from this study have implications for both athletic performance and rehabilitation. With athletic performance, these results imply that free ankle motion is an important
mechanical condition to allow higher reaction forces to be produced at the pedal, especially during the propulsive downstroke phase. The concept of "ankling," although not
explicitly studied here, makes sense. The ankle undergoes motion excursions throughout
one crank cycle with maximum plantar flexion occumng during the end of the downstroke
and maximum dorsiflexion occurring during the end of the upstroke. Our results show
that maintaining the ankle in a neutral position with a locked brace reduces the resultant
forces throughout the crank cycle, especially those imposed by maximum plantar flexion
during the power phase (i.e., at the end of the downstroke). Therefore, the results from this
study support the belief that cyclists can optimize pedaling performance by using an ankling
strategy.
In clinical rehabilitation settings, ankle impairments after musculoskeletal injury
and/or neurological insult can result in limited ankle motion, even after the injury has
healed. Bicycle ergometers, instrumented to measure pedal reaction forces, are frequently
used in research but are rarely used in clinical settings, yet they can provide valuable
information for both the clinician and patient during rehabilitation of ankle impairments
(Broker & Gregor, 1990;Newmiller et al., 1988). For example, Brown et al. (1994) showed
that real-time force feedback from each pedal can be used to correct asymmetries in force
production found in persons with poststroke hemiplegia. Our study also suggests that
feedback about ankle range of motion (e.g., provided by an electrogoniometer) might also
provide valuable information about pedaling performance and could be used to correct
dynamic ankle motion deficiencies that can occur when a person avoids using an injured
ankle. For example, in a person with a unilateral ankle impairment, pedaling an ergometer
Pedal Reaction Forces
345
might reinforce increased asymmetrical patterns of movement occurring from the impairment unless those patterns are recognized and corrected through feedback and changes in
pedaling strategies.
Ankle constraints may also be advantageous during rehabilitation. The ERGYS
powered leg cycle actually uses braces to constrain ankle joint movement for individuals
with spinal cord injury. Using certain stimulation parameters, Schutte, Rodgers, Zajac,
and Glaser (1993) designed a study to investigate the joint movements necessary to power
the legs. They found that by locking the ankles, thereby reducing the degrees of control
required by the stimulating device, they could forego stimulation of the ankle muscles and
optimize the cardiovascular benefits of this exercise for this patient population. Therefore,
clinicians can use a brace, locked or unlocked, at the ankle as a tool for either retraining
loss of dynamic control of ankle motion or providing extra support for the foot/pedal
interface during pedaling.
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Acknowledgments
We thank Steven A. Kautz, PhD, for help with analysis and interpretation of the data; James
Anderson and Douglas Schwandt, MSME, for technical assistance; and Kevin McGill, PhD, and
Felix E. Zajac, PhD, for their constructive comments about the manuscript.
This study was supported by funds from the Department of Veterans Affairs, Rehabilitation
Research and Development Division, and by a Career Development Grant from the American Association of University Women.