Quantification of the Effect of Dorsi-/
Plantarflexibility of Ankle-Foot Orthoses on
Hemiplegic Gait: A Preliminary Report
Sumiko Yamamoto, PhD
Masahiko Ebina, CPO
Shigeru Kubo
Hideo Kawai, RPT
Takeo Hayashi, CPO
Mitsuo Iwasaki, CPO
Toshio Kubota, MD
Shinji Miyazaki, PhD
ABSTRACT
A series of studies has been done to establish an objective and quantitative
method to determine the optimum
dorsi-/plantarfiexibility of an ankle-foot
orthosis (AFO) for various patients.
Introduction
Ankle-foot orthoses (AFOs) are commonly used in clinical practice to supplement the gait disabilities of hemiplegic patients secondary to cerebrovascular accidents and brain injuries.
However, few biomechanical research
studies have clarified how this improvement of gait is achieved in relation to AFOs' mechanical properties.
Factors involved in the selection and
prescription of AFOs include dorsi-/
plantarflexibility, initial setting of the
dorsi-/plantar angle and the length of
the sole. Among these factors, dorsi-/
plantarflexibility exerts the most significant effect on patients' gait.
Flexibility usually is controlled by selecting a proper trim line around the
malleolar aspect of the AFO. Often a
less-flexible AFO is prescribed because
of irreversibility of the trimming process and greater endurance, which generally results in so-called overbracing.
Biomechanical analyses of AFOs
have been made by Chowaniec et al.
and by Lehmann et al., but they have
not investigated the quantitative effect
of dorsi-/plantarfiexibility (1-4). Both
studies also limited measurement to a
single step although it is well-known
that the step-to-step variation of hemiplegic gait is large. Thus, their data are
difficult to interpret in a statistical
sense.
The authors have begun a series of
studies to establish an objective and
quantitative method to determine the
optimum dorsi-/plantarfiexibility of an
AFO for each patient. As the first step,
Yamamoto et al. have measured the
flexibility of various types of plastic
AFOs when fitted to the human limb
(5). Miyazaki et al. have made a mathematical model of gait with an AFO,
showing that the total ankle joint moment due to floor reaction forces
equals the sum of the moment generated by the muscles around the ankle
plus the corrective moment generated
by the AFO due to its deformation.
They then developed a system that
measured the total ankle joint moment
due to floor reaction forces and the corrective moment generated by the
AFO, continuously and separately t
estimate the ankle joint moment generated by the muscles (6). Based 01
these achievements, the purpose of the
study was to demonstrate the feasibility of quantifying the effect of dorsiplantarflexibility on hemiplegic gait in
a statistical manner.
MethodMeasurement
Measurement was made on a level
floor of 20 m in length. The following
quantities were measured simultaneously:
- total ankle-joint moment due to
floor reaction force
- corrective moment generated by
an AFO
- temporal factors
- step length
- ankle angle and knee angle in the
sagittal plane
- EMGs of plantarflexors and of
dorsiflexors.
The total ankle-joint moment was
measured by a device developed by Miyazaki et al. (6). This device uses a 2.5mm capacitive force transducer attached to the sole of the footwear that
measures the total moment within an
accuracy of +/- 3N-m Temporal factors
were extracted from the vertical foot
forces measured by the capacitive
transducer.
Because of the need to change the
dorsi-/plantarfiexibility of the AFO for
a short time during the experiment and
measure its corrective moment, a specially designed AFO was used (6). Figure 1
shows the experimental AFO,
which consists of a plastic shoe, modified medial and lateral double-Klenzak
joint, and a metal support. In each
Klenzak joint, two coil springs are
housed, and the dorsi-/plantarflexibility of the AFO can be controlled by
changing the stiffness of the springs. In
the study, soft, medium and hard
springs were used and are denoted as
spring No. 1, No. 2 and No. 3, respectively.
At the time of measurement, the
Klenzak joints were adjusted so their
centers were aligned to the center of
the ankle joint. The ankle joint angle
was then measured by a potentiometer
attached to the center of the lateral
Klenzak joint. This angle was used to
retrieve the corrective moment of the
AFO from the ankle.
The plastic shoe (PLS), made by cutting a conventional shoehorn-type
AFO at the ankle, can be removed
from the experimental AFO by releasing the screw bolts of the stirrup plate
connecting the Klenzak joints and the
PLS. The disassembled PLS was used
as experimental footwear to simulate
the patients' normal shoe. The medial
and lateral walls of the PLS were
trimmed to 25-30 mm so that the PLS
does not provide excessive medio-lateral stability. The subjects wore similar
PLSs on their sound feet during the
experiment. The experimental AFO,
including the PLS, weighed 1,060 g -
much more than the shoehorn-type
AFOs usually worn but comparable to
the weight of conventional double-metal-bar AFOs.
The step length was measured by a
simple device developed by Miyazaki
(7). The accuracy of the device is +/- 10
mm. The angle of the affected side in
the sagittal plane was measured by the
potentiometer attached to the experimental AFO when the subject walked
with the AFO, and was measured by a
potentiometer-type uni-axial goniometer when the subject walked with the
isolated PLS. The knee angle of the
affected side in the sagittal plane was
also measured by a potentiometer-type
uni-axis goniometers in both conditions.
EMGs of plantarflexors and of dorsiflexors of the affected side were measured by pairs of disposable surface
electrodes, electrically rectified and
low-pass filtered. Whole-body gait pattern was recorded from the lateral side
by a VTR for general reference.
Subjects
Two hemiplegic patients volunteered
as subjects. Subject A was a 64-yearold male left-hemiplegic patient who
suffered from cerebral infarction. He
was graded as stage III in the Brunnstrom scale. The measurement was
made seven months after the subject's
accident. He used a cane and usually
wore a shoehorn-type AFO. He is 166
cm tall, and weighs 62 kg.
Subject B was a 38-year-old male
right-hemiplegic patient who had suffered a brain injury. He also was graded as stage III. Subject B was measured
13 months after his accident and did
not use a cane but usually wore a shoehorn-type AFO. He is 167 cm tall and
weighs 64 kg.
Procedure
Measurements were first made with
the isolated PLS. The subjects were
asked to walk at a comfortable speed.
Initial and last portions of gait were
discarded, and the steady state gait was
measured for a continuous 30 seconds
in one path and sampled at 100 Hz.
Data recording the two paths was
stored for analysis. Before measurement, the subjects were allowed to
take a trial walk of the two paths to
become accustomed to the walkway
and the footwear.
The sequence was repeated with the
experimental AFO using different flexibilities: coil spring No. 1 and No. 2 for
subject A, and No. 1, No. 2 and No. 3
For subject B. To minimize fatigue, the
subjects rested between sessions.
Statistical Analysis
The following 24 quantitative parameters were extracted from raw data for
each step or cycle and were analyzed
for variance. One-way allocation was
assumed, and the flexibility of the
AFO was treated as the variation factor.
Temporal factors were cycle time,
average walking speed, stance phase of
the affected side, single support phase
of the affected side, double support
phase from the sound side to the affected side, and double support phase from
the affected side to the sound side.
Spatial factors were step length from
the affected side to the sound side and
step length from the sound side to the
affected side.
Joint angles (all for the affected side)
were maximum plantarfiexion during
stance and maximum dorsiflexion during stance, maximum plantarfiexion
during swing for the ankle joint, maximum extension during stance, maximum flexion during swing for the knee
joint.
Ankle joint moment: 1) maximum total ankle joint moment in the dorsiflex
direction where negative value indicates the subject and the AFO generate moment so the ankle joint is dorsiflexed, 2) maximum muscle moment in
dorsiflex direction, 3) maximum total
ankle joint moment in plantarflex direction, 4) maximum muscle moment
in plantarflex direction (all for the affected side), 5) maximum total ankle
joint moment in dorsiflex direction, 6)
maximum total joint moment in plantarflex direction (both for the sound
side).
Integrated EMG (all for the affected
side): 1) integrated value of EMG
(IEMG) of dorsiflexors during stance,
2) IEMG of dorsiflexors during swing,
3) IEMG of plantarflexors during
stance, 4) IEMG of plantarflexors during swing.
Results
Figure 2a
shows the relationship between the ankle joint angle and the corrective moment of the experimental AFO assembled with three coil springs: No. 1, No. 2 and No. 3. The
horizontal axis represents the ankle angle and the vertical axis the corrective
moment. A steeper curve indicates
lower flexibility.
Figure 2b
shows the relationship between the ankle joint angle and the corrective moment of three typical plastic
AFOs (5). The broken loop shows the
flexibility of a hemispiral AFO. The
two solid loops show the characteristics
of shoehorn-type AFOs. The steeper
loop represents a very rigid shoehorn-type AFO and the middle loop a flexible shoehorn-type AFO. Note that the
experimental AFO with coil spring No.
1, No. 2 and No. 3 roughly corresponds
to a hemispiral AFO, a flexible shoehorn-type AFO and a very rigid shoehorn-type AFO, respectively.
Figure 3
shows the typical waveforms of the ankle joint moment. In
each graph, the ankle joint moments of
the affected side for five consecutive
steps are superimposed. The horizontal axis represents time in seconds, and
its origin is the timing of foot contact
for each step. The vertical axis represents the ankle joint moment; the solid
line shows the total ankle joint moment
due to floor reaction forces, and the
broken line shows the corrective moment generated by the AFO. The results for subject A are shown in the
upper row.
The left graph shows the ankle joint
moment when the subject walked with
the isolated PLS. The middle and the
right graphs show the results when the
subject walked with the experimental
AFO assembled with coil springs No. 1
and No. 2, respectively. Similarly, the
results for subject B are shown in the
lower row. It is important to note the
magnitude of the total ankle joint moment. However, this small correcting
moment did generate noticeable differences in the pattern and timing of total
ankle moment.
The fact that a slight change in corrective moment resulted in a marked
difference in gait pattern is more clearly shown by the results of statistical
analysis of variance. In Figure 4
, Figure 5
, Figure 6
, Figure 7
, and Figure 8
, the
open circle and the filled triangle represent the results of subject A and subject B, respectively. In each graph, the
horizontal axis represents the level of
the variation factor in one-way allocation, that is, the flexibility of the AFO
(0 indicates that no AFO was used,
i.e., isolated PLS). This condition is
assumed to be the extreme case in
which the flexibility of the postulate
AFO is infinitive. Levels 1, 2 and
indicate the AFOs assembled with coil
springs No. 1, No. 2 and No. 3, respectively. As the number increases, the
flexibility decreases.
Results are plotted only when the
differences in the main effect are statistically significant at a confidence level
of .01. The vertical axis represents the
mean at each level. The left vertical
scale applies to the results of subject A
and the right vertical scale to those of
subject B. The number of samples at
each level was 25 to 27 for subject A
and 26 or 27 for subject B.
For subject A, plantarfiexion during
swing decreased monotonously as flexibility decreased (see Figure 6
). Since
one of the major functions of the AFO
is to prevent excessive plantarfiexion
during swing to assure better limb
clearance, degree of flexibility is important. In this sense, the least flexible
AFO is the best AFO.
However, such an AFO affects other
aspects of gait. It tends to cause forward thrust of the shank, causing knee
collapse after heel contact. It also
causes back knee in the late stage of
stance phase. Therefore, the best or
optimum flexibility is a trade-off between various factors.
From this point of view, it is important that all other parameters present
either maximum or minimum flexibility level 1, and that most can be interpreted as a sign of better gait function.
For example, the minimum cycle time
means fastest cadence. Combined with
the maximum step length from the affected side to the sound side and from
the sound side to the affected side, it
results in maximum walking speed (see
Figure 4
and Figure 5
). It is evident that the
optimum level of flexibility for subject
A in this situation is 1. It is interesting
to note that muscle moment in plantarflex direction increases when the subject wears the AFO (see Figure 7
).
For subject B, there is no clear tendency governing all parameters. It may
be said that the optimum level of flexibility for this subject is 2 since the walking speed and the step lengths are at
maximum at this level (see Figures 4
and Figure 5
). It must be noted that the maximum extension during stance, i.e.,
back knee, also occurs at this level (see
Figure 6
). This subject walks at maximum speed with the AFO assembled
with coil spring No. 2 at a risk of future
complications at the knee. Another interesting observation is that IEMG of
the plantarflexors and the muscle moment in plantarflex direction vary considerably with flexibility (see Figure 7
and Figure 8
).
Discussion
The gait patterns of two hemiplegic patients were quantified and subjected to
statistical analysis with the aid of recently developed continuous measurement techniques. Preliminary results demonstrated the basic feasibility
of quantifying the effect of dorsi-/plantarfiexibility of an AFO on the gait,
thus determining the optimum flexibility for each patient.
Present techniques involve several
clinical problems. The first problem is
that the experimental AFO employed
in this study is too heavy and may not
simulate plastic AFOs worn daily. To
solve this problem, a new, lighterweight experimental AFO is being designed.
The second problem is the time and
labor of the measurement. It took at
least 90 minutes to complete the measurement for four levels of flexibility.
Most of this time was consumed attaching measuring instruments to the subject and changing the coil springs.
Time may be saved by selecting the
most critical parameters to reduce the
number of measuring instruments. In
the experimental AFO, the coil springs
housed in the Klenzak joints will be
replaced by larger coil springs and
placed outside the joints to ease handling. These improvements will help
reduce the physical and mental burden
on the subject and the number of people necessary for the measurement.
Quantitative measurement and subsequent statistical analysis may not always
provide sufficient information to determine optimum flexibility. A good example was given in the case of subject B. He
walked at maximum speed when using
the AFO assembled with coil springs
No. 2, but he also presented the maximum back knee at this level.
Careful examination of the waveform of the knee angle and the images
recorded in the VTR revealed that this
subject had an unstable knee, and the
back knee occurred only at the last
stage of the stance phase for a rather
short duration. The moderate back
knee occurred at flexibility levels of 0,
1 and 3. Intense clinical study must be
made to assess how hazardous this
maximum back knee at flexibility level 2 is the subject and to determine optimum flexibility.
When analyzing variance in this
study, one-way allocation was assumed, and flexibility was treated as
the only variation factor. Theoretically, a two-way allocation model should
have been used in which flexibility and
weight of the AFO were designated as
variation factors. However, no essential difference in interpretation would
have resulted because weight had the
same effect on the mean of each parameter at flexibility levels 1, 2 and 3.
In the case of flexibility level 0, i.e.,
PLS, it had a different effect since the
isolated PLS is much lighter than the
assembled AFO.
This neglect of the lightweight factor
created bias in favor of the PLS because the gait with PLS would have
been otherwise impeded by the additional weight. Still, the optimum occurred at flexibility level 1 or 2. Thus,
the conclusion would have been the
same.
The observation that maximum muscle moment in plantarflex direction
varied significantly with the use of the
AFO and its flexibility in both subjects
is contrary to the common understanding of the function of AFOs. It is generally believed that the active ankle moments generated by paralyzed plantarflexors do not change considerably regardless of the use of AFOs because
the active ankle moment is uniquely
determined by the degree of paralysis
of the muscles.
One of the main functions of AFOs
is to supplement the decreased plantarflex moment late in the stance phase.
However, the present measurement
clearly demonstrated that the corrective supplementary moment generated
by the AFO is very small compared
with the total ankle moment and the
net muscle moment did change significantly (see Figure 3
). In particular,
plantarflex muscle moment increased
when subject A walked with the AFO.
In the case of subject B, maximum
plantarflex muscle moment occurred
when he walked without the AFO.
However, considering the weight factor mentioned above, the maximum
plantarflex moment might have occurred with the AFO at a flexibility
level of 2 in this subject if the AFO had
been much lighter.
Although we have not yet investigated the underlying mechanism(s) of this
new observation, we have two working
hypotheses. One possibility is that the
subject activated maximum plantarflexor force because lateral stability
was augmented by the AFO. The subject could have been relieved from a
fear of lateral instability. The second
possibility is that the activity level of
the plantarflexors is determined by the
dynamic matching between the mechanical factors and human pathological factors although the range of activity is basically limited by the underlying
paralysis of the muscles.
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