Use the labels in the right column to find what you want. Or you can go thru them one by one, there are only 34,264 posts. Searching is done in the search box in upper left corner. I blog on anything to do with stroke. DO NOT DO ANYTHING SUGGESTED HERE AS I AM NOT MEDICALLY TRAINED, YOUR DOCTOR IS, LISTEN TO THEM. BUT I BET THEY DON'T KNOW HOW TO GET YOU 100% RECOVERED. I DON'T EITHER BUT HAVE PLENTY OF QUESTIONS FOR YOUR DOCTOR TO ANSWER.
Changing stroke rehab and research worldwide now.Time is Brain!trillions and trillions of neuronsthatDIEeach day because there areNOeffective hyperacute therapies besides tPA(only 12% effective). I have 523 posts on hyperacute therapy, enough for researchers to spend decades proving them out. These are my personal ideas and blog on stroke rehabilitation and stroke research. Do not attempt any of these without checking with your medical provider. Unless you join me in agitating, when you need these therapies they won't be there.
What this blog is for:
My blog is not to help survivors recover, it is to have the 10 million yearly stroke survivors light fires underneath their doctors, stroke hospitals and stroke researchers to get stroke solved. 100% recovery. The stroke medical world is completely failing at that goal, they don't even have it as a goal. Shortly after getting out of the hospital and getting NO information on the process or protocols of stroke rehabilitation and recovery I started searching on the internet and found that no other survivor received useful information. This is an attempt to cover all stroke rehabilitation information that should be readily available to survivors so they can talk with informed knowledge to their medical staff. It lays out what needs to be done to get stroke survivors closer to 100% recovery. It's quite disgusting that this information is not available from every stroke association and doctors group.
Showing posts with label Stiff-Knee Gait. Show all posts
Showing posts with label Stiff-Knee Gait. Show all posts
Background and Purpose—
Stiff-knee gait is defined as reduced knee flexion during the swing
phase. It is accompanied by frontal plane compensatory movements (eg,
circumduction and hip hiking) typically thought to result from reduced
toe clearance. As such, we examined if knee flexion assistance before
foot-off would reduce exaggerated frontal plane movements in people with
stiff-knee gait after stroke.
Methods— We
used a robotic knee orthosis to assist knee flexion torque during the
preswing phase in 9 chronic stroke subjects with stiff-knee gait on a
treadmill and compared peak knee flexion, hip abduction, and pelvic
obliquity angles with 5 nondisabled control subjects.
Results—
Maximum knee flexion angle significantly increased in both groups, but
instead of reducing gait compensations, hip abduction significantly
increased during assistance in stroke subjects by 2.5°, whereas no
change was observed in nondisabled control subjects. No change in pelvic
obliquity was observed in either group.
Conclusions—
Hip abduction increased when stroke subjects received assistive knee
flexion torque at foot-off. These findings are in direct contrast to the
traditional belief that pelvic obliquity combined with hip abduction is
a compensatory mechanism to facilitate foot clearance during swing.
Because no evidence suggested a voluntary mechanism for this behavior,
we argue that these results were most likely a reflection of an altered
motor template occurring after stroke.
Stiff-knee gait (SKG), defined as reduced and delayed knee flexion during the swing phase,1
is a common disability in people with stroke. Individuals with SKG
often exhibit exaggerated frontal plane movements such as hip
circumduction and hip hiking.2 These frontal plane movements are energetically demanding3 and may lead to chronic joint pain.4 There is believed to be a cause-and-effect relationship between these two sets of movements.1
Reduced knee flexion decreases an individual’s ability to clear the
foot during the swing phase. In turn, subjects make compensatory frontal
plane movements that facilitate toe clearance. Although this
compensatory theory is commonly accepted, there are no quantitative in
vivo studies examining the interaction between the swing phase knee
flexion angle and proximal frontal plane kinematic behaviors in
individuals with poststroke hemiparesis. If a kinematic connectivity
indeed exists, it would indicate that the observed frontal plane
behaviors may in part be a response to mechanical constraints imposed at
the knee.
Alternatively, exaggerated frontal plane motions may be
due to abnormal motor control. Recent studies in patients with stroke
have found abnormal across-joint torque activation coupling, or
synergies, between hip and knee activation. For example, several studies
have found evidence of abnormal coupling of knee flexion/extension and
hip abduction/adduction activity in both reflex5 and voluntary6,7
muscle activity. These findings suggest that improving swing phase knee
flexion angle may not resolve exaggerated frontal plane movements.
Accordingly,
we sought to examine the effect of assistive knee flexion torque during
preswing on frontal plane gait compensations. To achieve this goal, we
developed a lightweight, backdrivable actuator capable of selectively
applying knee flexion torque during the preswing phase of gait without
imposing any significant mechanical constraints on the subject.8
To assess if the observed frontal plane behaviors during the assisted
steps were a result of a mechanically coupled response to the
“perturbation” or due to motor control compensation to the assistance,
catch trials (no assist) were introduced at random.9
If kinematic changes persist during the no assist trials, it would
suggest the presence of an acute motor adaptation induced by knee
flexion assistance.10
Thus, to examine the potential effect of anticipatory motor commands to
the repeated assistance, kinematic comparisons between the assisted and
unanticipated steps with no assistance (catch trials) were conducted.
Quantitative metrics of gait stability11
were used to evaluate the potential effect that the knee flexion
assistance may have had on locomotor stability in both the stroke and
healthy participants. Knowledge gleaned from this study will improve our
understanding of the relationship between knee impairments and the
proximal 3-dimensional movements in people with SKG poststroke.
Specifically, this study will help identify whether abnormal frontal
plane behaviors are a response to local knee impairment or a
manifestation of an abnormal multisegmental motor program. We argue that
identifying these aberrant biomechanical interactions will help guide
future clinical investigations and the development of assistive
technology for individuals with stroke.
Successful
walking requires the execution of the pre-swing biomechanical tasks of
body propulsion and leg swing initiation, which are often impaired
post-stroke. While excess rectus femoris activity during swing is often
associated with low knee flexion, previous work has suggested that
deficits in propulsion and leg swing initiation may also contribute. The
purpose of this study was to determine underlying causes of propulsion,
leg swing initiation and knee flexion deficits in pre-swing and their
link to stiff knee gait in individuals post-stroke.
(My problem is that most of my pre-motor cortex is dead and nothing here addresses that. Do you people ever consider that you're supposed to be solving stroke problems for everyone?)
Methods
Musculoskeletal
models and forward dynamic simulations were developed for individuals
post-stroke (n = 15) and healthy participants (n = 5). Linear
regressions were used to evaluate the relationships between peak knee
flexion, braking and propulsion symmetry, and individual muscle
contributions to braking, propulsion, knee flexion in pre-swing, and leg
swing initiation.
Results
Four
out of fifteen of individuals post-stroke had higher plantarflexor
contributions to propulsion and seven out of fifteen had higher vasti
contributions to braking on their paretic leg relative to their
nonparetic leg. Higher gastrocnemius contributions to propulsion
predicted paretic propulsion symmetry (p = 0.005) while soleus
contributions did not. Higher vasti contributions to braking in
pre-swing predicted lower knee flexion (p = 0.022). The rectus femoris
had minimal contributions to lower knee flexion acceleration in
pre-swing compared to contributions from the vasti. However, for some
individuals with low knee flexion, during pre-swing the rectus femoris
absorbed more power and the iliopsoas contributed less power to the
paretic leg. Total musculotendon work done on the paretic leg in
pre-swing did not predict knee flexion during swing.
Conclusions
These
results emphasize the multiple causes of propulsion asymmetry in
individuals post-stroke, including low plantarflexor contributions to
propulsion, increased vasti contributions to braking and reliance on
compensatory mechanisms. The results also show that the rectus femoris
is not a major contributor to knee flexion in pre-swing, but absorbs
more power from the paretic leg in pre-swing in some individuals with
stiff knee gait. These results highlight the need to identify individual
causes of propulsion and knee flexion deficits to design more effective
rehabilitation strategies.
Background
Over
795,000 people in the United States experience a stroke each year and
over half of individuals post-stroke over age 65 have reduced mobility [1].
Regaining walking ability is an important goal of rehabilitation as
walking speed is a critical predictor of long-term health [2]
and individuals post-stroke who achieve limited or full community
walking speeds report an overall higher quality of life than those who
remain household ambulators [3].
Successful walking requires the execution of the critical pre-swing
biomechanical subtasks of body propulsion and leg swing initiation,
which are often impaired post-stroke [4, 5] and may influence swing phase knee flexion [2, 6,7,8].
For
example, modeling studies having identified knee flexion velocity at
toe-off as the primary contributor to peak knee flexion during swing [9] and low push-off acceleration has also been linked to stiff knee gait [2]. Moreover, impaired knee flexion is often attributed to rectus femoris activity [10, 11]
and a modeling study showed that eliminating rectus femoris activity in
pre-swing was more effective than eliminating rectus femoris activity
in early swing for improving knee flexion [7].
Decreased gastrocnemius activity may also contribute to stiff knee
gait, as increased gastrocnemius contributions to pre-swing knee flexion
were observed after gait retraining [8].
However, a representative individual post-stroke with a limited
community walking speed had lower iliopsoas contributions to leg swing
initiation in pre-swing but similar contributions from the gastrocnemius
compared to a healthy control [12].
Thus, while the potential of lower extremity muscles to increase or
decrease knee flexion velocity in late stance has been documented [13], it is unknown which muscles most affect pre-swing knee flexion velocity in individuals post-stroke.
Braking and propulsion deficits are also common in individuals post-stroke [4], and in addition to predicting slower walking speeds [14, 15], may contribute to stiff knee gait. The plantarflexors are primary contributors to propulsion [16, 17]. Decreased plantarflexor contributions to propulsion have been observed in individuals post-stroke [12, 18, 19], which could occur due to both muscle activation deficits [20, 21] and altered muscle and Achilles tendon properties [22,23,24]. The gastrocnemius is an important contributor to both propulsion and leg swing initiation [17],
and thus low propulsion may be related to knee flexion deficits.
Stimulating the plantarflexors in pre-swing increases peak knee flexion
for individuals post-stroke [25]. However, the total propulsive force did not predict knee flexion [6]. On average, individuals with impaired plantarflexor coordination do not have lower propulsion, but rather greater braking [26] likely due to co-activation of the plantarflexors and vasti muscles, which are primary contributors to braking [9].
Knee flexion velocity at toe-off may be diminished by late braking
forces because muscles such as the vasti and rectus femoris that
contribute to braking also contribute to knee extension and oppose leg
swing initiation in late-stance [9]. In addition, late stance braking forces in individuals post-stroke predict less knee flexion during swing [6].
However, it is unknown if the relationships between braking, propulsion
and swing phase kinematics are causal or correlative. Due to the
characteristically high variability between individuals post-stroke,
there are a number of mechanisms that can cause both propulsion and knee
flexion deficits.(So you're justifying your failure to solve this because it is hard. Well, fuck you then. Hope you're OK with this when you're the 1 in 4 per WHO that has a stroke.)
Previous work has established the importance of
pre-swing conditions to achieving adequate swing phase knee flexion.
However, actual muscle contributions to propulsion, knee velocity and
leg-swing initiation in individuals post-stroke and their relationship
to swing-phase knee flexion has not been established. Thus, the
objectives of this study were to determine the underlying causes of
propulsion and braking deficits and identify muscle contributors to
pre-swing knee flexion acceleration and leg swing initiation in
individuals post-stroke with and without stiff knee gait. We
hypothesized that (1) braking and propulsion asymmetries would be caused
by both low plantarflexor contributions to propulsion and high vasti
contributions to braking, (2) vasti and plantarflexor contributions to
propulsion and braking in pre-swing would predict swing phase knee
flexion, (3) the rectus femoris would be a major contributor to knee
extension in pre-swing in individuals with stiff knee gait, and (4)
total musculotendon power delivered to the leg in pre-swing would
predict knee flexion during swing. The outcomes of this work will
highlight specific deficits in propulsion and leg swing initiation
post-stroke and their implications for swing phase knee flexion, which
will provide a basis for developing targeted rehabilitation strategies.
Background and Purpose—
Stiff-knee gait is defined as reduced knee flexion during the swing
phase. It is accompanied by frontal plane compensatory movements (eg,
circumduction and hip hiking) typically thought to result from reduced
toe clearance. As such, we examined if knee flexion assistance before
foot-off would reduce exaggerated frontal plane movements in people with
stiff-knee gait after stroke.
Methods— We
used a robotic knee orthosis to assist knee flexion torque during the
preswing phase in 9 chronic stroke subjects with stiff-knee gait on a
treadmill and compared peak knee flexion, hip abduction, and pelvic
obliquity angles with 5 nondisabled control subjects.
Results—
Maximum knee flexion angle significantly increased in both groups, but
instead of reducing gait compensations, hip abduction significantly
increased during assistance in stroke subjects by 2.5°, whereas no
change was observed in nondisabled control subjects. No change in pelvic
obliquity was observed in either group.
Conclusions—
Hip abduction increased when stroke subjects received assistive knee
flexion torque at foot-off. These findings are in direct contrast to the
traditional belief that pelvic obliquity combined with hip abduction is
a compensatory mechanism to facilitate foot clearance during swing.
Because no evidence suggested a voluntary mechanism for this behavior,
we argue that these results were most likely a reflection of an altered
motor template occurring after stroke.
Stiff-knee gait (SKG), defined as reduced and delayed knee flexion during the swing phase,1
is a common disability in people with stroke. Individuals with SKG
often exhibit exaggerated frontal plane movements such as hip
circumduction and hip hiking.2 These frontal plane movements are energetically demanding3 and may lead to chronic joint pain.4 There is believed to be a cause-and-effect relationship between these two sets of movements.1
Reduced knee flexion decreases an individual’s ability to clear the
foot during the swing phase. In turn, subjects make compensatory frontal
plane movements that facilitate toe clearance. Although this
compensatory theory is commonly accepted, there are no quantitative in
vivo studies examining the interaction between the swing phase knee
flexion angle and proximal frontal plane kinematic behaviors in
individuals with poststroke hemiparesis. If a kinematic connectivity
indeed exists, it would indicate that the observed frontal plane
behaviors may in part be a response to mechanical constraints imposed at
the knee.
Alternatively, exaggerated frontal plane motions may be
due to abnormal motor control. Recent studies in patients with stroke
have found abnormal across-joint torque activation coupling, or
synergies, between hip and knee activation. For example, several studies
have found evidence of abnormal coupling of knee flexion/extension and
hip abduction/adduction activity in both reflex5 and voluntary6,7
muscle activity. These findings suggest that improving swing phase knee
flexion angle may not resolve exaggerated frontal plane movements.
Accordingly,
we sought to examine the effect of assistive knee flexion torque during
preswing on frontal plane gait compensations. To achieve this goal, we
developed a lightweight, backdrivable actuator capable of selectively
applying knee flexion torque during the preswing phase of gait without
imposing any significant mechanical constraints on the subject.8
To assess if the observed frontal plane behaviors during the assisted
steps were a result of a mechanically coupled response to the
“perturbation” or due to motor control compensation to the assistance,
catch trials (no assist) were introduced at random.9
If kinematic changes persist during the no assist trials, it would
suggest the presence of an acute motor adaptation induced by knee
flexion assistance.10
Thus, to examine the potential effect of anticipatory motor commands to
the repeated assistance, kinematic comparisons between the assisted and
unanticipated steps with no assistance (catch trials) were conducted.
Quantitative metrics of gait stability11
were used to evaluate the potential effect that the knee flexion
assistance may have had on locomotor stability in both the stroke and
healthy participants. Knowledge gleaned from this study will improve our
understanding of the relationship between knee impairments and the
proximal 3-dimensional movements in people with SKG poststroke.
Specifically, this study will help identify whether abnormal frontal
plane behaviors are a response to local knee impairment or a
manifestation of an abnormal multisegmental motor program. We argue that
identifying these aberrant biomechanical interactions will help guide
future clinical investigations and the development of assistive
technology for individuals with stroke.
Stiff-Knee
gait (SKG) after stroke is often accompanied by decreased knee flexion
angle during the swing phase. The decreased knee flexion has been
hypothesized to originate from excessive quadriceps activation. However,
it is unclear whether hyperreflexia plays a role in this activation.
The goal of this study was to establish the relationship between
quadriceps hyperreflexia and knee flexion angle during walking in
post-stroke SKG.
Methods
The
rectus femoris (RF) H-reflex was recorded in 10 participants with
post-stroke SKG and 10 healthy controls during standing and walking at
the pre-swing phase. In order to attribute the pathological
neuromodulation to quadriceps muscle hyperreflexia and activation,
healthy individuals voluntarily increased quadriceps activity using
electromyographic (EMG) feedback during standing and pre-swing upon RF
H-reflex elicitation.
Results
We observed a negative correlation (R = − 0.92, p = 0.001)
between knee flexion angle and RF H-reflex amplitude in post-stroke
SKG. In contrast, H-reflex amplitude in healthy individuals in presence (R = 0.47, p = 0.23) or absence (R = − 0.17, p = 0.46)
of increased RF muscle activity was not correlated with knee flexion
angle. We observed a body position-dependent RF H-reflex modulation
between standing and walking in healthy individuals with voluntarily
increased RF activity (d = 2.86, p = 0.007), but such modulation was absent post-stroke (d = 0.73, p = 0.296).
Conclusions
RF
reflex modulation is impaired in post-stroke SKG. The strong
correlation between RF hyperreflexia and knee flexion angle indicates a
possible regulatory role of spinal reflex excitability in post-stroke
SKG. Interventions targeting quadriceps hyperreflexia could help
elucidate the causal role of hyperreflexia on knee joint function in
post-stroke SKG.
Introduction
Stiff-Knee gait (SKG) is one of the most common gait disabilities following stroke. SKG is defined as reduced knee flexion [30] during the swing phase. Those with SKG have joint pain [16], energy inefficiency due to compensatory motions [9, 36, 38] and increased risk of falls [3]. Post-stroke SKG has been attributed to overactivity of rectus femoris (RF) muscle [2, 13, 14] and decreased activity of ankle plantar flexors and iliopsoas that generate knee flexion moment [23, 29]. Quadriceps muscle overactivity is the most widely accepted cause of SKG [18, 33, 42].
To this end, Botulinum toxin (Botox) injections that block
acetylcholine release in the femoral nerve show modest improvements in
knee flexion [34, 35, 39],
suggesting that rectus femoris (RF) reflex excitability contributes to
SKG. However, the cause of excessive RF muscle activity is unclear. One
hypothesis is that increased quadriceps activation could be achieved
voluntarily to improve stability during the stance phase, but then fails
to relax during the pre-swing phase [18, 33].
RF overactivity could additionally be explained as spasticity in the
form of reflex hyperexcitability and lack of reciprocal inhibition.
Accumulating
evidence suggests that reduced knee flexion in post-stroke SKG depends
on pathological modulation of spinal reflex loops. For example, Lewek et
al. [27]
found increased quadriceps short-latency reflex excitability following
hip extension perturbations in post-stroke individuals. The degree of
hyperexcitability was correlated to knee flexion angle, suggesting that
altered involuntary responses could play a role in diminished knee
flexion during the swing phase of post-stroke SKG. Others found that hip
abduction perturbations elicited abnormally coordinated RF activation,
suggesting a role of abnormal reflex-mediated coordination in
post-stroke gait [12]. Using a custom robotic actuator [41] to perturb knee flexion during pre-swing in individuals with post-stroke SKG [40],
we observed a sharp knee extension velocity following initial increased
knee flexion angle during swing. However, no such reaction was found
during steps where the assistance was temporarily removed, or in a
baseline period before the assistance was applied, indicating that the
knee extension was induced by the robotic assistance. The knee extension
was preceded by increased RF electromyographic (EMG) activity within a
short latency following the perturbation, and further musculoskeletal
modeling analysis showed a correlation between RF fiber stretch velocity
and RF activity [1].
Taken together, this evidence points to RF reflex hyperexcitability
influencing knee flexion in post-stroke SKG. Representative altered
reflex pathways in post-stroke include reduced gait phase-dependent
modulation [8, 21] and changes in presynaptic inhibition [11]. However, their relation to SKG has yet to be determined.
The
objective of this study(wrong objective, nothing here helps survivors recover) was to characterize the relation of
hyperreflexia to knee flexion in post-stroke SKG. We investigated reflex
excitability via the monosynaptic H-reflex with well-established
neuronal pathways [20]. H-reflexes are a reliable and consistent probe in identifying the altered neuronal pathways [28].
We examined the modulation of monosynaptic RF H-reflexes during
standing and walking in people post-stroke compared to healthy control
subjects in order to establish the extent to which reflex modulation is
related to the knee flexion angle after stroke. In order to determine
the possibility that hyperreflexia is a byproduct of improper timing of
RF activity, we compared our results to healthy controls with
volitionally up-regulated RF activity during pre-swing. We hypothesized
that RF H-reflex hyperexcitability during standing and walking in
pre-swing phase in people with post-stroke is associated with decreased
knee flexion during swing phase in SKG compared to healthy controls.
Characterization of the role of hyperreflexia in SKG and other
neuromuscular disorders could help to design targeted treatments for
hyperreflexia including but not limited to operant H-reflex conditioning
training [43].
Ok, you have described a problem and discussed it a lot, but no solution. So useless. My lower leg swing is very poor due to spasticity. Cure my spasticity and I will walk normally in no time.
Background and Purpose—
Stiff-knee gait is defined as reduced knee flexion during the swing
phase. It is accompanied by frontal plane compensatory movements (eg,
circumduction and hip hiking) typically thought to result from reduced
toe clearance. As such, we examined if knee flexion assistance before
foot-off would reduce exaggerated frontal plane movements in people with
stiff-knee gait after stroke.
Methods— We
used a robotic knee orthosis to assist knee flexion torque during the
preswing phase in 9 chronic stroke subjects with stiff-knee gait on a
treadmill and compared peak knee flexion, hip abduction, and pelvic
obliquity angles with 5 nondisabled control subjects.
Results—
Maximum knee flexion angle significantly increased in both groups, but
instead of reducing gait compensations, hip abduction significantly
increased during assistance in stroke subjects by 2.5°, whereas no
change was observed in nondisabled control subjects. No change in pelvic
obliquity was observed in either group.
Conclusions—
Hip abduction increased when stroke subjects received assistive knee
flexion torque at foot-off. These findings are in direct contrast to the
traditional belief that pelvic obliquity combined with hip abduction is
a compensatory mechanism to facilitate foot clearance during swing.
Because no evidence suggested a voluntary mechanism for this behavior,
we argue that these results were most likely a reflection of an altered
motor template occurring after stroke.
Stiff-knee gait (SKG), defined as reduced and delayed knee flexion during the swing phase,1
is a common disability in people with stroke. Individuals with SKG
often exhibit exaggerated frontal plane movements such as hip
circumduction and hip hiking.2 These frontal plane movements are energetically demanding3 and may lead to chronic joint pain.4 There is believed to be a cause-and-effect relationship between these two sets of movements.1
Reduced knee flexion decreases an individual’s ability to clear the
foot during the swing phase. In turn, subjects make compensatory frontal
plane movements that facilitate toe clearance. Although this
compensatory theory is commonly accepted, there are no quantitative in
vivo studies examining the interaction between the swing phase knee
flexion angle and proximal frontal plane kinematic behaviors in
individuals with poststroke hemiparesis. If a kinematic connectivity
indeed exists, it would indicate that the observed frontal plane
behaviors may in part be a response to mechanical constraints imposed at
the knee.
Alternatively, exaggerated frontal plane motions may be
due to abnormal motor control. Recent studies in patients with stroke
have found abnormal across-joint torque activation coupling, or
synergies, between hip and knee activation. For example, several studies
have found evidence of abnormal coupling of knee flexion/extension and
hip abduction/adduction activity in both reflex5 and voluntary6,7
muscle activity. These findings suggest that improving swing phase knee
flexion angle may not resolve exaggerated frontal plane movements.
Accordingly,
we sought to examine the effect of assistive knee flexion torque during
preswing on frontal plane gait compensations. To achieve this goal, we
developed a lightweight, backdrivable actuator capable of selectively
applying knee flexion torque during the preswing phase of gait without
imposing any significant mechanical constraints on the subject.8
To assess if the observed frontal plane behaviors during the assisted
steps were a result of a mechanically coupled response to the
“perturbation” or due to motor control compensation to the assistance,
catch trials (no assist) were introduced at random.9
If kinematic changes persist during the no assist trials, it would
suggest the presence of an acute motor adaptation induced by knee
flexion assistance.10
Thus, to examine the potential effect of anticipatory motor commands to
the repeated assistance, kinematic comparisons between the assisted and
unanticipated steps with no assistance (catch trials) were conducted.
Quantitative metrics of gait stability11
were used to evaluate the potential effect that the knee flexion
assistance may have had on locomotor stability in both the stroke and
healthy participants. Knowledge gleaned from this study will improve our
understanding of the relationship between knee impairments and the
proximal 3-dimensional movements in people with SKG poststroke.
Specifically, this study will help identify whether abnormal frontal
plane behaviors are a response to local knee impairment or a
manifestation of an abnormal multisegmental motor program. We argue that
identifying these aberrant biomechanical interactions will help guide
future clinical investigations and the development of assistive
technology for individuals with stroke.
Methods
Nine hemiparetic persons with stroke (Table)
gave written informed consent according to the guidelines approved by
Northwestern University and Department of Veterans Affairs Institutional
Review Boards.
Table. Data for 9 Hemiparetic Stroke Subjects*
Subject No.
Age, Years
G
W,kg
Postyrs
Meds
AFO
Speed,m/s
Modified Ashworth
Knee (U/I),Degrees/Degrees
Hip,Degrees
Pelvis,Degrees
KE
KF
*All
subjects who wore an AFO used a hinged version. “Speed” refers to each
subject’s self-selected gait speed during overground walking. Modified
Ashworth scores are for ankle knee extensors (KE) and knee flexors (KF)
based on a 0 to 5 scale. “Knee” refers to the range of motion
(unimpaired and impaired sides) of the knee during gait without the
orthosis. “Hip” and “Pelvis” refer to the difference of maximum hip
abduction and pelvic obliquity of the impaired limb and unimpaired
limbs, respectively. Subject 7 did not have her Ashworth scores
measured.
G
indicates gender; W, weight; Post, years poststroke; Meds, taking
Baclofen; AFO, use of an ankle–foot orthosis; M, male; F, female; N, no;
Y, yes.
1
49
M
59
1
N
N
0.80
0
0
71/16
1
13
2
18
M
52
2
Y
Y
0.76
0
0
54/32
9
7
3
59
F
58
16
N
N
0.53
3
1
53/19
−16
18
4
38
M
91
1
Y
Y
0.72
1
0
60/26
1
6
5
51
F
75
3
Y
Y
0.92
1
0
52/15
6
12
6
40
F
82
3
Y
N
1.1
3
0
54/34
−3
11
7
31
F
59
3
N
Y
0.74
54/32
6
6
8
67
M
54
7
N
N
0.38
2
3
64/30
−1
5
9
43
M
85
29
N
N
1.0
0
1
54/38
4
9
Mean
44
68
7
0.78
57/27
1
10
SD
15
15
9
0.23
6/8
7
4
Although SKG after stroke has clinically been recognized as reduced and delayed knee flexion angle during the swing phase,1
a clear quantitative metric for the severity in SKG is lacking. Hence,
we included subjects whose knee range of motion was at least 16° less on
the affected limb than the unimpaired side. Subjects also had to be
able to walk continuously for 20 minutes at 0.55 m/s. All participants
were left-sided hemiparetics. Exclusion criteria were (1) knee
hyperextension during the stance phase (genu recurvatum); and (2)
inability to provide informed consent. Five nondisabled control subjects
also completed the protocol. Some of the control data have been
previously published.10
A powered knee orthosis known as the SERKA (Series Elastic Remote Knee Actuator) was used in the study.8
The SERKA consists of a modified off-the-shelf anterior cruciate
ligament knee brace that can be adjusted to fit comfortably about the
paretic knee for each individual subject. A servomotor located away from
the user controls the excursion of a sheathed (Bowden) cable
transmission. The cable pulls on a capstan coupled to the brace by a
compliant torsional spring controlling for spring deflection and, thus,
torque (Figure 1).
The device is capable of exerting large torques (28 N · m) in 40 ms
with a maximum torque of 41 N · m. Because the motor is remotely
located, the weight of the device on the leg is 1.2 kg (approximately
10% of limb mass). The unpowered device exerted no appreciable
resistance (torque <1 N · m) to the subject’s walking ability. During
treadmill walking, no visual cues were given to the subjects. In
addition, the actuators did not provide auditory cues detectable over
the background noise of the treadmill.
Figure 1. The
experimental setup consisted of a subject walking on a split-belt force
treadmill used to measure individual ground reaction forces. Kinematics
was measured using motion capture cameras and reflective markers
attached to both legs. Each subject wore a powered knee brace to provide
knee flexion torque.
Lower
limb kinematics were collected at 100 Hz using an 8-camera video system
(Motion Analysis, Santa Rosa, Calif). Thirty markers were placed
bilaterally on the lower limbs and pelvis.12
Both
the control and stroke subjects were instructed to walk on an
instrumented split-belt ADAL treadmill (Tecmachine, Andrez Boutheon,
France) at 0.55 m/s, approximately the average speed for people with
SKG.13
Subjects were instructed to walk as naturally as possible and allowed
to hold onto the treadmill handrails. Subjects were also instructed to
try to walk with 1 foot on either tread but to prioritize walking
naturally. Each subject donned a harness to protect from falls but not
support weight. An emergency stop switch was available to both the
subject and experimenter (Figure 1). Subjects were permitted to wear their ankle–foot orthoses.
The
treadmill vertical ground reaction forces were used to calculate the
maxima and minima of center of pressure of both the paretic and
nonparetic feet to identify gait events.14
For control subjects, the powered knee orthosis exerted a flexion
torque that began at contralateral heel strike and ended at swing phase
initiation. However, stroke subjects had a longer double-support period;
therefore, torque onset began 250 ms before paretic swing phase
initiation (predicted from the timing of the previous step). This timing
more closely matches the double-support time of healthy gait. The ramp
time of the torque pulse was set at 170 ms. Preliminary data showed that
the changes in ramp time did not affect outcome measures; hence, the
ramp time was chosen to maintain an acceptable level of comfort for the
stroke participants. Analog data, including force plate data and applied
torque, was acquired at 1 kHz.
An initial torque calibration test
was conducted during treadmill walking (45 steps) to determine the
appropriate amplitude of knee flexion assistance for each of the
participants. We applied knee flexion torque during preswing to random
individual steps during the torque calibration test. Knee flexion torque
began at 10 N · m and increased by approximately 3 N · m during each
application to a maximum of 35 N · m. We selected the torque assistance
amplitude to be used for the data acquisition test from post hoc torque
calibration test data that corresponded to the greatest peak knee
flexion during swing or 60° of knee flexion, whichever was lower.
The data acquisition test consisted of 610 steps, lasting approximately 16 minutes (Figure 2).
During the first 50 steps (baseline), no forces were exerted on the
subject. The last 10 steps of this phase were used as baseline data.
During the next 560 steps (assistance), subjects received a knee flexion
torque assistance every step as described previously. We collected 15
trials in 10-step increments interspersed throughout this phase. Also
interspersed were 4 nonconsecutive trials with no torque assistance
during a single step cycle (catch trials). Subjects were instructed to
report any perceived changes in their gait patterns during the
experimental paradigm to obtain a qualitative assessment of performance.
Figure 2. The
experimental protocol consisted of 3 treadmill-walking stages. The
torque calibration test determined the magnitude of the torque used
during assistance. During this stage, a range of knee flexion torques
was applied to determine an appropriate level of assistance for each
subject. After the baseline stage (no assistance), the assistance stage
provided constant, open-loop flexion torque assistance during the
preswing phase of each step. There were 4 catch trials without
assistance in this stage.
All
data were imported into Matlab (Natick, Mass) with marker data and
torque data time-synchronized. Kinematic data were filtered with a
fifth-order low-pass Butterworth filter at 8 Hz. Each stride was
separated and normalized at left foot strike determined using the
maximum forward displacement of the toe marker. Pelvic obliquity was
determined by calculating the rotation matrix between the coordinate
system of the pelvis15
and world coordinates followed by Euler angle decomposition. Because
markers on the affected thigh were placed on the brace, therefore not
directly on the thigh, thigh segment position was determined by a
coordinate system, including the projected hip center from the pelvis
and medial and lateral knee center from the shank. Hip abduction was
calculated based on the Euler angle decomposition between this projected
thigh coordinate system and the pelvis. Knee flexion angle was
determined from the rotation matrix between the projected thigh and
shank coordinates. Absolute angles were calculated from standing data.
Center of mass movement was estimated from the centroid of the sacrum
and left and right anterior–superior iliac spine markers.16 Kinematics were linearly interpolated to match the difference in sampling frequency with torque data.
The
main outcomes consist of the peak kinematics of the paretic side during
the swing phase; specifically, maximum knee flexion and simultaneously
with peak knee flexion, hip abduction, and pelvic obliquity. The
kinematics during assistance was compared with baseline and catch
trials. Any potential anticipation or adaptation would be found in catch
trials. Thus, we used repeated-measures analysis of variance (ANOVA)
with 3 levels: baseline, assistance (subject mean), and catch trials
(subject mean) followed by Tukey-Kramer post hoc testing. A significance
level of α=0.05 was selected. We also used a 2-sample t test to compare these parameters between stroke and control data.
To
assess the potential effect of loss of balance on the observed
kinematic adaptations, the extrapolated center of mass was calculated.
Comparing the extremes of this outcome, which comprised of lateral
center of mass position and velocity, with lateral center of pressure
resulted in a margin that, when negative, showed an instability or lean.11
Another indirect measure of lean can be calculated through the sum of
the lateral forces on the nonparetic foot. Both of these outcomes were
compared using the same 3-level repeated-measures ANOVA mentioned
previously.
Results
Peak
kinematics without the orthosis shows a higher incidence of exaggerated
pelvic obliquity but no correlation with peak knee flexion angle. When
stroke subjects walked without the orthosis, peak knee flexion in the
affected limb was 30.4±12.0° (mean±SD) less than the healthy limb (t test; P<0.0001). In addition, asymmetry was observed in pelvic obliquity (10±4.0°, t test; P<0.001) between the healthy and affected sides but not in peak hip abduction (1.0±7.0°, t test; P=0.75).
Pelvic obliquity and hip abduction asymmetry were not significantly
correlated with peak knee flexion angle asymmetry (Pearson, P=0.34 and 0.59, respectively).
Both
groups increased knee and hip flexion angle in response to the
assistance. Selected kinematic parameters are shown for representative
subjects of the control and stroke groups in Figure 3 and Figure 4,
respectively. Stroke subjects significantly increased peak knee flexion
angle by 13.8±11.4° when receiving the supplemental knee flexion torque
from the powered knee orthosis (repeated-measures ANOVA; F[2,8]=9.21, P=0.002), and control subjects’ knee flexion angle increased significantly by 12.4±3.3° above baseline (ANOVA; F[2,4]=26.0, P<0.001).
In both groups, knee flexion angle during knee flexion torque
assistance was significantly greater than during catch trials (Tukey
honestly significant difference [THSD]; P<0.05). There was no significant difference in knee flexion between the first and last steps of assistance (t test, P>0.05) for either group (Figure 5). There was no significant difference between the change in peak knee flexion of the control and stroke groups (t test, P>0.05).
Figure 3. Kinematics
from a representative control subject. Each trace is an average and 95%
CI for 4 steps. Knee flexion torque increases peak knee flexion, hip
flexion angle, and toe clearance, but has no apparent effect on frontal
plane characteristics.
Figure 4. Kinematics
from a representative stroke subject. Each trace is an average and 95%
CI for 10 steps. At peak toe height, knee flexion, and hip flexion
angle, when compensations are needed the least, more hip abduction
occurs.
Figure 5. Key
kinematic outcome measures show increase knee flexion of both control
(left) and stroke (right) subjects, but a significant increase in hip
abduction only for stroke subjects. Horizontal dashed line at zero
represents baseline and is used for reference.
Control subjects significantly increased hip flexion angle during assistance by 9.1±4.4° (ANOVA; F[2,4]=15.7, P=0.002) and significantly more than during catch trials (THSD; P<0.05). Stroke subjects also increased hip flexion 9.9±8.8° (ANOVA; F[2,8]=7.42, P=0.005), but not significantly higher than during catch trials (THSD; P>0.05).
Results
from the repeated-measures ANOVA revealed small but significant shifts
in hip abduction in response to assistance in the stroke group,
2.31±1.15° above baseline (ANOVA; F[2,8]=21.0, P<0.001). No significant difference was detected for control subjects (ANOVA; F[2,4]=3.54, P=0.079). In stroke subjects, hip abduction increased significantly compared with baseline and catch trials (THSD; P<0.05). Neither group exhibited a significant difference from initial to final steps of assistance (t test, P>0.05; Figure 5). There were no significant differences between catch trials and baseline for stroke subjects or for control subjects (THSD; P>0.05). Stroke subjects’ change in hip abduction was significantly higher than control subjects’ (t test, P<0.05).
Peak pelvic obliquity significantly changed during this experiment for stroke (ANOVA; F[2,8]=5.00, P=0.021), but not in control subjects (ANOVA; F[2,4]=1.96, P=0.203).
In the stroke group, catch trials were significantly greater than both
baseline (1.24±1.32°) and assistance (1.22±1.68°; THSD; P<0.05), but no difference was found between assistance and baseline (THSD; P>0.05). Neither group exhibited a significant difference between first and last steps of assistance (t test, P>0.05; Figure 5).
Finally,
our statistical examination of the stability margins showed no
significant change compared with baseline in either the stroke subjects
(−7±19 cm, ANOVA; F[2,8]=1.04, P=0.37) or the control subjects (10±16 cm, ANOVA; F[2,4]=0.95, P=0.42)
across conditions (baseline, catch, and assistance). The lateral force
impulse on the right foot also showed no change from baseline in either
the stroke (−25±67 N · s, ANOVA; F[2,8]=0.66, P=0.53) or control (−58±43 N · s, ANOVA; F[2,4]=3.82, P=0.069)
groups. After the initial adaptation trials, subjects reported no
perception of lack of balance during the knee assistance trials.
Discussion
In
this study, we sought to examine the role of preswing knee flexion on
frontal plane movement in SKG after stroke. As expected, we found an
increase in knee flexion angle during swing when subjects received knee
flexion torque assistance during preswing. Surprisingly, our findings
also indicated that hip abduction increased when stroke subjects
received assistive knee flexion torque at foot-off despite improved
ability to clear the foot, whereas control subjects had no such
behavior. Our data further indicated that these changes were neither a
reflection of assistance-induced instability in the frontal plane nor
leaning. These findings are in direct contrast to the traditional belief
that hip abduction is solely a compensation to clear the foot during
swing. Because no evidence suggested a voluntary mechanism for
exaggerated abduction, we argue that these results were most likely a
reflection of an altered motor template occurring after stroke.
One
may argue that the increased hip abduction during assistance observed
in stroke subjects is a strategy to enhance locomotor stability or as a
reaction to the imposed knee flexion assistance. No significant changes
in stability margins or leaning forces for either group were observed as
a result of the assistance. Additionally, no subjects reported any
perception of imbalance by midtraining. One may also argue that
exaggerated abduction is an anticipatory gesture, but no such effect was
observed in catch trials. These findings suggest that the observed
kinematic changes are not due to imbalance, but may be a manifestation
of a more intrinsic motor behavior mediated by the stroke.
The
peak hip flexion angle of the control group was greater during
assistance than during the catch trials. It is likely that these
differences may be due in part to an increase in the shank-to-thigh
intersegmental coupling mediated by the applied sagittal plane torque at
the knee.17
This mechanical effect was muted in the stroke group. Stroke subjects’
peak hip flexion angle calculated during the assist trials was similar
to those calculated during the unanticipated catch trials. This hip
angle invariance provides indirect evidence of an abnormal
neuromechanical coupling between the hip and knee sagittal plane
mechanics in the stroke group.18
Our
3-dimensional analysis of the proximal kinematics indicated that hip
abduction increased significantly compared with baseline and catch
trials in the stroke group, whereas no frontal plane changes were
observed in healthy control subjects. The presence of abnormal muscle
synergies provides one possible explanation for the observed knee
flexion-mediated increase in hip abduction. A recent examination of
lower limb volitional synergies found a coupling between voluntary hip
abduction and knee flexion in patients with stroke in a static position
imitating paretic foot-off.7
Hence, it is possible that the greater hip abduction observed in this
study may have been a reflection of such across-joint synergies observed
under static conditions, a knee flexion coupled with hip abduction. It
may also be possible that the exaggerated frontal plane movement was due
to abnormal knee–hip reflex couplings, similar to those found
previously in a seated posture in the same population.5
To explore this, a thorough examination of lower limb muscle activity
associated with the assistance is warranted. Delineating the
differential effects of these multisegmental mechanisms to the observed
behaviors is likely to improve clinical treatment of gait dysfunction
after stroke and will provide guidance for future clinical and basic
science investigations.
Musculoskeletal models are increasingly used to investigate underlying biomechanical factors in gait abnormalities after stroke19;
such abnormalities are characterized by excessive frontal plane
movements. However, most of the existing model-based analyses of such
pathologies have thus far been limited to the sagittal plane mechanics.18–20
Given the significant 3-dimensional interactions presented in this
study, we argue that special attention should be given to the use of
generic musculoskeletal models in the study of pathological gait.
Specifically, our data indicated that an increase in knee flexion
velocity preswing resulted in an increase in the peak knee flexion angle
during swing, a finding that is consistent with musculoskeletal
modeling predictions.21,22
However, although sagittal plane behaviors were consistent across
groups, 3-dimensional kinematic differences between the control and
stroke groups persisted in response to the externally applied knee
flexion torque. One could argue that these differences are potentially a
manifestation of poststroke neural constraints.23
Taken together, inclusion of such neural constraints in future
computational studies may increase the efficacy of musculoskeletal
modeling for the study of pathological gait such as in stroke.
Our
goal in this study was to examine the kinematic connectivity between
knee sagittal plane kinematic impairments and frontal plane movements
proximal to the knee. In this context, subjects were instructed to “walk
as naturally as possible.” Our data suggest that abnormal frontal plane
gait patterns persisted regardless of the imposed knee flexion
assistance. It remains to be seen if, with additional kinematic feedback
or instruction,24
subjects would be able to volitionally reduce these aberrant frontal
plane movements during gait. Thus, combining knee flexion assistance
with supplemental strategies that reduce frontal plane gait
compensations may improve walking efficiency in those with SKG after
stroke.
Finally, our findings indicate that facilitation of knee
flexion did not lead to the restoration of symmetrical gait behavior
mediated by reduced hip hiking and normal hip abduction/adduction gait
patterns. To this end, the current study provides a basic examination of
recent developments in the assistive technology targeting the design of
adaptive knee flexion tools to restore asymmetrical gait.
