Download Biomechanical changes elicited by an anterior cruciate ligament

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Clinical Biomechanics xxx (2003) xxx–xxx
www.elsevier.com/locate/clinbiomech
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Michael A. Hunt a, David J. Sanderson
Helene Moffet c, J. Timothy Inglis
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c
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School of Human Kinetics, The University of British Columbia, 210-6081 University Blvd, Vancouver, BC, Canada
b
ICORD––International Collaboration on Repair Discoveries, Canada
Department of Rehabilitation, Faculty of Medicine, Laval University and CIRRIS Research Center, Quebec City, QC, Canada
Received 9 October 2002; accepted 26 February 2003
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9 Abstract
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Objectives. To identify any changes to lower limb biomechanics during steady rate cycling as a result of an anterior cruciate
ligament deficiency.
Design. Comparative study in which healthy and anterior cruciate ligament injured individuals underwent biomechanical analysis
during stationary cycling.
Background. Individuals with an anterior cruciate ligament deficiency often exhibit reductions in the magnitude of quadriceps
muscle activity and subsequent knee joint extensor moments during walking. It is not known whether these compensations are
present during cycling, an exercise frequently used to retrain anterior cruciate ligament injured individuals.
Methods. Ten healthy and 10 unilateral anterior cruciate ligament deficient individuals participated. All participants were required to cycle for approximately 30 s at each of six different cycling intensities while lower limb EMG, kinetics, and kinematics were
collected bilaterally. Before riding, participants performed submaximal isometric contractions to generate normalizing data.
Results. In addition to reduced quadriceps activation and net knee joint extensor moments, the anterior cruciate ligament deficient limbs exhibited decreases in linear impulse of the resultant pedal force, knee joint flexor moments, hip and ankle extensor
moments, and muscle activity from gluteus maximus. These decreases were counteracted by an increase in output from the anterior
cruciate ligament intact limb.
Conclusion. Anterior cruciate ligament injured individuals exhibited a limb attenuation strategy during cycling activities.
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a,b,*
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Biomechanical changes elicited by an anterior cruciate
ligament deficiency during steady rate cycling
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Relevance This study reports lower limb kinetic and electromyographic data from anterior cruciate ligament deficient individuals
26 during stationary cycling, and shows that these individuals exhibit a limb attenuation strategy on the very leg that is undergoing
27 rehabilitation.
28 Ó 2003 Published by Elsevier Science Ltd.
29 Keywords: Cycling; Anterior cruciate ligament; Quadriceps avoidance
30 1. Introduction
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The term ‘‘quadriceps avoidance’’ was coined by
Berchuck et al. (1990), who reported a net knee joint
flexor moment during the mid-stance phase of gait (10–
30% of the gait cycle) in anterior cruciate ligament
(ACL) deficient limbs. This is in contrast to many
publications that report that in non-injured individuals
*
Corresponding author. Address: School of Human Kinetics, The
University of British Columbia, 210-6081 University Blvd, Vancouver,
BC, Canada V6T 1Z1.
E-mail address: [email protected] (D.J. Sanderson).
0268-0033/03/$ - see front matter Ó 2003 Published by Elsevier Science Ltd.
doi:10.1016/S0268-0033(03)00046-9
the net knee joint moment during this phase is consistently extensor (e.g. Winter, 1983). Although electromyographic (EMG) data were not collected in the study,
Berchuck et al. (1990) proposed that the net flexor
moments were the result of a preferential decrease in the
magnitude of activation of the quadriceps muscle group.
That is, as a result of the injury to the ACL, the individual selectively de-recruited the knee extensor muscles,
perhaps as a protection mechanism.
Two studies have reported significant reductions in
the magnitude of the net knee joint extensor moment
(Devita et al., 1997; Chmielewski et al., 2001), while
Roberts et al. (1999) reported no change in the magni-
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2.1. Participants
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Twenty participants volunteered for the study and
gave their informed consent in accordance with the
University Ethical Review Board. Ten (five males and
five females; mean age: 25.1 (SD 4.7) years) individuals
with a unilateral deficiency of the ACL and 10 age and
gender-matched healthy uninjured control individuals
(mean age: 25.9 (SD 3.4) years) participated. Body
proportions of the ACL group (68.9 (SD 10.8) kg body
mass, 90.1 (SD 6.0) cm leg length) were similar to the
controls (64.4 (SD 9.9) kg body mass, 86.2 (SD 3.1) cm
leg length). ACL-injured individuals presented with a
unilateral rupture of the ACL, as diagnosed by positive
Lachman and anterior drawer tests, at a mean time of
10.7 (SD 13.9) months post-injury, were capable of full
knee joint range of motion, and were pain-free during
everyday activity.
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2.2. Experimental task
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Participants performed six randomized bouts of stationary cycling at intensities comprised of two cadences
(60 and 90 rpm) and three power outputs generally used
in an ACL injury rehabilitation program (75, 125, and
175 W). EMG, bi-directional pedal reaction force, and
kinematic data were collected bilaterally for a total of 18
s after the participants reached the correct cadence and
were able to maintain it (5%) at the pre-chosen power
output. The total time required for each trial did not
exceed 2 min and adequate rest was given between trials,
negating any effects of fatigue for the riders.
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2.3. Instrumentation
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Participants rode on a standard racing bicycle
mounted on a Velodyne Trainer (Schwinn, Chicago, IL,
USA) which enabled manipulation of power output,
while cadence was monitored using a Cateye cyclocomputer (Cateye Co., Boulder, CO, USA) attached to
the bicycle. Seat height was manipulated such that the
vertical distance from the seat to the pedal at bottomdead-center (BDC) was equal to the vertical distance
from the subjectÕs greater trochanter to the floor (Nordeen-Snyder, 1977). Crank position data were collected
using a photoelectric cell positioned at top-dead-center
(TDC) for the left pedal, which gave an analog pulse
when triggered. Pedal angle (with respect to the crank)
data were collected continuously using a Dynapar digital encoder (Danaher Controls, Gurnee, IL, USA) attached to each pedal.
EMG data were collected from the muscle belly of
rectus femoris, vastus lateralis, biceps femoris, and
semitendinosis using bipolar surface electrodes (Thera-
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2. Methods
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tude of the net knee joint moments in a similar design.
Evidence of a preferential decrease in quadriceps muscle
activation comes from EMG studies reporting significantly reduced quadriceps muscle activity in ACL-deficient individuals during gait (Limbird et al., 1988; van
Lent et al., 1994). However, these reported decreases in
quadriceps muscle activity did not always occur during
the mid-stance phase associated with the reduced net
knee joint extensor moments. Because calculated joint
moments are net moments, they reflect the activity of
agonist and antagonist muscles. Reported reduced net
knee joint extensor moments may also occur as a consequence of increased flexor muscle activity. In fact,
increases in hamstring muscle activation have also been
reported in ACL-deficient individuals during gait
(Limbird et al., 1988). However, the timing during which
these increases occurred did not always coincide with the
reductions in the net knee joint extensor moments. It
would seem that these variations in joint moments and
muscle activation reflect a more complex interaction of
muscle strength, integrity, and neuromuscular control.
One difficulty in exploring this phenomenon is the
many degrees of freedom of movement associated with
walking. Consequently, we considered an alternate
model. Stationary cycling is an activity that is similar to
walking in that both walking and cycling require––in
alternating phases of propulsion and recovery––output
from both limbs for forward propulsion. Similarity in
the patterns of joint kinematics, joint moments of force
(Winter, 1983; Gregor et al., 1985), and patterns of
muscle activation for the quadriceps, hamstrings, and
gluteus maximus (Winter, 1984; Mohr et al., 1981) lend
strength to the efficacy of using cycling as a means to
explore quadriceps avoidance. Coincidentally, stationary cycling is a common exercise used in rehabilitation
programs for an ACL rupture as it has been proposed to
increase cardiovascular endurance, knee joint range of
motion, and lower limb muscular strength (McLeod and
Blackburn, 1980). However, no studies have been published that have reported the biomechanical characteristics of ACL-deficient individuals during cycling
exercise.
The purpose of the present study was to examine the
kinetic, EMG, and kinematic patterns in ACL-deficient
individuals during stationary cycling. We propose that
this model will provide an effective means to explore the
notion of quadriceps avoidance. Further, it was postulated that ‘‘quadriceps avoidance’’, specifically, reduced
net knee joint extensor moments and reduced quadriceps muscle activity, would be present during stationary
cycling.
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2.5. Statistical analysis
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complete cycle. Subject averages were computed as the
average over 15 consecutive cycles. Group averages were
then calculated for each of four limbs: control individuals right limb (CON R), control individuals left limb
(CON L), ACL individuals intact (ACL I), and ACL
individuals deficient limb (ACL D).
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3. Results
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All variables were analyzed separately using a 4
(limb) 2 (cadence) 3 (power output) mutifactorial
analysis of variance (A N O V A ) with repeated measures on
the last two factors. Significant F -ratios were further
analyzed with a Tukey HSD post-hoc test and differences were classified based on a significance level of 0.05.
3.1. Muscle EMG
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Average normalized iEMG values for a single cycle
are shown in Fig. 1a–e. Significant differences were exhibited between limbs for rectus femoris (F3;36 ¼ 21:87,
P < 0:001), vastus lateralis (F3;36 ¼ 14:53, P < 0:001),
and gluteus maximus (F3;36 ¼ 5:65, P ¼ 0:003), while no
differences existed for the two hamstring muscles: biceps
femoris (F3;36 ¼ 1:84, P ¼ 0:157) and semitendinosis
(F3;36 ¼ 1:13, P ¼ 0:348). Post-hoc analysis revealed that
compared to control limbs, ACL intact limbs exhibited
greater muscle activation for the rectus femoris and
vastus lateralis muscles, while ACL-deficient limbs
showed less activation for the rectus femoris, vastus
lateralis, and gluteus maximus muscles. Although the
magnitudes of EMG activation levels were different
between limbs for certain muscles, the general pattern
and timing of the activation were similar between limbs
and across cycling intensities (Fig. 2a–e).
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3.2. Pedal forces
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Linear impulses, calculated from the resultant forces
data, are summarized in Fig. 3. A significant limb effect
was observed (F3;36 ¼ 15:87, P < 0:001), and post-hoc
analysis revealed that for all intensities the impulse of
the two control limbs were less than the ACL intact
limbs (left limb: P ¼ 0:042, right limb: P ¼ 0:013),
greater than the ACL-deficient limbs (left limb:
P ¼ 0:001, right limb: P ¼ 0:005), while not being significantly different from each other.
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3.3. Joint moments
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peutics Unlimited, Iowa City, IA, USA). EMG data
were also collected from the muscle belly of gluteus
maximus using pre-gelled surface electrodes (Red Dot
2259, 3M Company, Borken, Germany) and an Octopus
AMT-8 amplifier (Bortec, Calgary, AB, Canada). Prior
to electrode application, the designated area was shaved
and cleansed with alcohol to reduce electrical impedance. Raw EMG data were collected at 600 Hz using a
Data Translation 3010 analog-to-digital (A/D) converter (Data Translation, Marlboro, MA, USA) and a
Peak Performance Technologies Data Acquisition System (Peak Performance Inc., Denver, CO, USA). Prior
to testing, participants were required to perform a submaximal isometric contraction with each muscle to
generate normalizing data. With the participants seated,
a strap was placed around the lower limb at the height
of the lateral malleolus, and participants were instructed
to maintain a force of 75 N with visual feedback as
measured by a force transducer (Artech S-Beam, Riverside, CA, USA).
Each bicycle pedal was instrumented with two force
transducers (Kistler Instruments, Winterthur, Switzerland) capable of measuring normal (Fz ) and shear (Fy )
forces applied to the top of the pedal as has been previously used in the lab (Sanderson et al., 2000). Kinetic
data were also collected using the Data Translation A/D
converter and Peak acquisition system.
Kinematic data were collected at a sampling rate of
60 Hz using the peak acquisition system and two cameras, positioned 3 m from the sagittal plane of the bicycle. Reflective markers were placed bilaterally on the
skin overlying the greater trochanter, lateral femoral
condyle, lateral malleolus, lateral aspect of the calcaneus, and the lateral aspect of the fifth metatarsal. Reflective markers were also placed over the lateral aspect
of the force transducers to denote the center of pressure
of force application. Raw coordinate data were acquired
by digitizing each frame for each of the 18 second data
collection periods.
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190 2.4. Data reduction
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All raw data (rectified EMG, kinetic, raw coordinate
data) were filtered using a 4th-order low-pass Butterworth filter with a cut-off frequency of 4 Hz. Raw coordinate data were used to calculate segmental center of
mass (COM) coordinates, joint angles, as well as segmental linear and angular velocities and accelerations to
be combined with kinetic data to calculate joint moments of force using conventional inverse dynamics
principles. Linear impulse was calculated from the resultant force vector applied to each pedal for a complete
cycle (TDC to subsequent TDC). A ratio of filtered task
EMG to filtered normalization EMG was calculated for
each data point and then used to integrate the magnitude of the signal (iEMG) for each muscle for each
Peak moments were analyzed for the hip, knee, and 248
ankle joints and are reported in Table 1. Significant limb 249
effects occurred for peak ankle extensor (F3;36 ¼ 25:03, 250
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Fig. 1. Group mean (SD) integrated EMG values for the 175 W and 60 rpm condition. Values shown are the ratio of EMG collected during riding to
EMG collected during the normalizing contractions for rectus femoris (RF), vastus lateralis (VL), biceps femoris (BF), semitendinosis (ST), and
gluteus maximus (GM) for each of the four limbs: CON L, control left; CON R, control right; ACL I, ACL intact; ACL D, ACL deficient.
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P < 0:001), knee flexor (F3;36 ¼ 20:22, P < 0:001), and
hip extensor (F3;36 ¼ 21:86, P < 0:001) moments. Posthoc analysis revealed that in all cases where significant
limb effects were reported, no asymmetry was reported
between control limbs, while significant asymmetries occurred between limbs of ACL-deficient individuals. All
asymmetries exhibited in ACL injured individuals were
the result of the peak joint moment in the ACL intact limb
being significantly larger than the ACL-deficient limb.
260 3.4. Joint kinematics
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Although there was a tendency for ACL individuals
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and less ankle plantarflexion, no significant differences
were observed between limbs for maximum thigh extension (F3;36 ¼ 1:47, P ¼ 0:239), maximum knee extension (F3;36 ¼ 1:51, P ¼ 0:229), or maximum ankle
plantarflexion (F3;36 ¼ 0:731, P ¼ 0:540).
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4. Discussion
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Results from the present study provide evidence for a
quadriceps avoidance strategy in ACL-deficient limbs
during cycling similar to that reported during walking
(Berchuck et al., 1990). ACL-deficient limbs exhibited a
measurable reduction (although not statistically signifi-
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Fig. 2. Group mean EMG ensemble averages with respect to crank angle for control (solid lines), ACL intact (dotted line), and ACL-deficient
(dashed line) limbs for the 175 W and 60 rpm condition. Values on the vertical axis are the normalized data. Graphs correspond to rectus femoris (A),
vastus lateralis (B), biceps femoris (C), semitendinosis (D), and gluteus maximus (E) muscles.
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cant) in the magnitude of the peak net knee joint extensor moment. This occurred due to reductions in
rectus femoris and vastus lateralis muscle activity and
no change in hamstrings (biceps femoris and semitendinosis) activity. However, in contrast to walking, where
increased hip or ankle extensor moments have been reported in ACL-deficient limbs to compensate for the
reduced knee joint extensor moment (Berchuck et al.,
1990; Chmielewski et al., 2001; Roberts et al., 1999),
ACL-deficient individuals in the present study tended to
reduce output across the entire injured limb.
This apparent limb attenuation was manifested in
significant decreases in the linear impulse of the resultant pedal force, maximum extensor moments at the hip
and ankle and flexor moment at the knee, and magnitude of muscle activation in rectus femoris, vastus lateralis, and gluteus maximus in ACL-deficient limbs
compared to uninjured limbs. The differences exhibited
were not due a change in movement path since the
ranges of motion about all three lower limb joints were
similar between injured and uninjured limbs. Further,
the timing and general pattern of muscle activation was
similar for all limbs (Fig. 2), suggesting that a similar
motor program was used for all limbs.
It has been suggested that changes in knee joint biomechanics in ACL-deficient individuals during gait
occur as a result of decreased stability within the knee
joint due to increased anterior tibial translation (Ber-
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Impulse (Ns)
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Cycling Condition
Table 1
Group mean (SD) peak extensor and flexor moments
125/60
88.5 (13.9)
92.9 (15.3)
106.3 (12.7)
76.1 (15.2)
114.4 (20.0)
116.4 (19.9)
135.3 (20.2)
90.5 (12.9)
Knee
CON L
CON R
ACL I
ACL D
8.8 (3.6)
8.8 (4.5)
11.1 (3.4)
7.6 (4.7)
7.9 (6.6)
7.4 (6.4)
13.4 (5.6)
8.5 (6.3)
Ankle
CON L
CON R
ACL I
ACL D
23.9
23.6
28.2
19.4
(3.3)
(3.3)
(3.1)
(3.8)
33.6
32.0
37.3
25.5
Flexor
Hip
CON L
CON R
ACL I
ACL D
15.2
18.1
21.1
12.6
(8.6)
(11.0)
(6.8)
(9.6)
Knee
CON L
CON R
ACL I
ACL D
42.7
42.6
44.4
26.3
Ankle
CON L
CON R
ACL I
ACL D
-2.1
-2.0
-2.8
-2.5
175/60
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75/60
Extensor
Hip
CON L
CON R
ACL I
ACL D
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Fig. 3. Group mean (SD) data for linear impulse for each of the six conditions. Symbols indicate control left (r), control right (j), ACL intact (N),
and ACL-deficient () limbs.
125/90
175/90
72.7 (14.7)
83.7 (14.2)
105.7 (15.5)
70.3 (10.6)
108.5 (16.1)
115.5 (14.1)
129.3 (17.5)
84.1 (23.6)
138.6
139.5
153.3
109.1
12.5 (8.0)
16.4 (8.0)
13.6 (6.9)
9.2 (8.6)
9.3 (3.2)
10.6 (3.4)
15.3 (5.8)
11.4 (3.2)
13.8
14.9
17.9
12.1
(5.9)
(6.1)
(7.1)
(5.1)
17.2
20.3
17.6
13.7
(5.5)
(6.7)
(6.3)
(6.0)
(4.9)
(4.0)
(4.3)
(3.9)
40.3
39.2
45.1
30.6
(4.0)
(3.3)
(7.6)
(7.1)
24.2
24.5
32.6
23.8
(3.2)
(4.2)
(5.5)
(4.4)
31.7
30.7
35.3
25.1
(3.5)
(1.7)
(3.9)
(3.8)
40.1
38.3
41.4
30.7
(5.7)
(5.6)
(8.8)
(5.7)
17.3
20.4
23.2
20.6
(11.0)
(10.3)
(6.6)
(13.9)
27.1
32.1
31.2
31.7
(15.3)
(17.6)
(15.3)
(11.1)
16.1
27.1
27.5
20.0
(9.3)
(11.2)
(9.9)
(6.6)
24.5
25.7
25.5
24.9
(16.1)
(14.1)
(12.0)
(6.6)
27.3
35.7
29.8
20.5
(19.1)
(12.7)
(11.1)
(9.3)
(9.6)
(9.2)
(12.6)
(6.4)
49.9
49.8
57.4
38.9
(9.0)
(6.2)
(11.8)
(9.0)
59.3
58.4
56.9
33.8
(8.1)
(7.6)
(13.2)
(10.2)
31.8
34.4
43.6
27.3
(5.4)
(6.0)
(11.9)
(6.5)
42.2
43.6
51.6
32.8
(5.9)
(6.0)
(9.4)
(8.1)
59.4
60.1
64.8
39.5
(11.1)
(9.7)
(9.0)
(7.4)
(0.9)
(1.4)
(2.1)
(1.3)
)0.9
)1.2
)1.0
)1.5
(0.7)
(1.2)
(2.0)
(2.4)
0.8
0.3
0.8
0.4
)1.3
)0.1
)1.3
)0.2
(0.8)
(1.4)
(2.0)
(1.4)
)1.7
)1.5
)1.9
)0.9
(0.9)
(2.6)
(2.2)
(1.5)
)1.1 (1.0)
0.1 (1.0)
)0.8 (2.3)
)0.1 (2.1)
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141.8 (18.1)
141.6 (20.2)
154.9 (9.1)
99.4 (9.3)
(1.2)
(2.2)
(2.7)
(2.7)
Values are in N m.
denotes a significantly decreased peak joint moment in ACL-deficient limbs compared to ACL intact limbs (P < 0:05).
*
(16.0)
(16.5)
(16.3)
(15.5)
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5. Conclusions
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Results from the present study indicate that a pattern
similar to the quadriceps avoidance strategy observed in
gait (decreased quadriceps muscle activation, decreased
knee joint extensor moment) also occurs during stationary cycling. However, during cycling injured individuals exhibited a limb attenuation of the entire ACLdeficient limb that was manifested in decreases in extensor moments at the hip, knee, and ankle, as well as
muscle activity from the quadriceps and gluteus maximus muscles. We believe that this limb attenuation is
unique to cycling and only possible because an increase
in output from the contralateral limb can maintain
output at the crank in light of decreased output from the
injured limb. Understanding this phenomenon is important because this limb attenuation strategy occurs on
the very leg that is undergoing rehabilitation.
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Winter (1980).
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similar to all other limbs. This large asymmetry between
limbs in the ACL group is in contrast to gait, where a
decrease in output from the ACL intact limbs––in order
to maintain symmetry with the decreased output from
the injured limb––has been reported (van Lent et al.,
1994). A decrease in output from both limbs in the ACL
group was impossible in this study as the force imparted
to the crank must have remained constant to maintain
the proper power output at each cadence.
One of the most important aspects of a rehabilitation
program is the formation of new motor programs in
order to compensate for the injury (Kvist and Gillquist,
2000). If an ACL-deficient individual learns a limb
avoidance strategy early in their rehab program this
strategy could somehow manifest itself to a lesser extent
in other movements. Based on the results from the
present study, there appears to be similar biomechanical
compensations in ACL-deficient limbs during stationary
cycling and walking suggesting that cycling may be an
effective model to examine quadriceps avoidance. Since
stationary cycling is an integral part of early rehabilitation for this injury, it is possible that the quadriceps
avoidance pattern observed during walking may be associated with compensations made early during the rehabilitation period, and stationary cycling in particular.
However, this is speculation and further research is required to examine this possible connection especially as
the goal of any rehabilitation strategy is to return the
limb to full function.
EC
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chuck et al., 1990; Devita et al., 1997; Ferber et al.,
2002). In order to decrease the tibial translation, it has
been argued that individuals with an ACL deficiency
reduce the quadriceps activation, which will act to pull
the tibia forward, while increasing hamstrings activation
(a known ACL synergist McNair et al., 1992; More
et al., 1993; Imran and OÕConnor, 1998; Liu and
Maitland, 2000), which will act to pull the tibia backward, at a time when anterior tibial translation is known
to be largest during gait (Marans et al., 1989; Lafortune
et al., 1992; Kvist and Gillquist, 2000). These strategies
will, therefore, result in a more flexed knee and a decreased knee joint extensor moment.
The amount of anterior tibial translation is not
known during cycling, but can be estimated based on
tibiofemoral shear force (Neptune and Kautz, 2000) and
ACL strain (Fleming et al., 1998) data. It is known that
both tibiofemoral shear force and ACL strain are largest
during the propulsive phase of cycling, which would
suggest that anterior tibial translation is also largest at
that time. It is also known that larger forces at the pedal
result in increased anterior translation (Fleming et al.,
2001). Therefore, a simple strategy to decrease the
amount of anterior tibial translation during stationary
cycling would be to decrease the magnitude of the force
exerted to the pedal by the ACL-deficient limb during
propulsion, which is what was found in this study.
If the extensor output at the knee joint is decreased
during walking in ACL-deficient individuals in order to
maintain tibial positioning, then compensations must be
made at the other joints to maintain an adequate support moment as well as forward propulsion (Winter,
1983). It has been shown that ACL-deficient individuals
increase the magnitude of the hip (Berchuck et al., 1990;
Roberts et al., 1999) or ankle (Chmielewski et al., 2001)
extensor moment during mid-stance to compensate for
the decreased extensor moment about the knee. This
prevents the individual from falling down and helps
maintain adequate propulsion from the injured limb.
That a compensation at the hip or ankle joint was not
observed in the present study is likely due to two fundamental differences between walking and cycling. First,
the weight of the body is supported by the bicycle seat.
Therefore, the maintenance of a support moment is not
necessary. Second, since both pedals are connected to
the crank, it is possible for a lack of propulsion from one
leg to be compensated for by an increase in propulsion
from the other limb.
In the present study, a limb attenuation in the ACLdeficient limbs, where output was decreased to limit the
amount of anterior tibial translation, was possible because the decrease in output from the ACL-deficient
limb was counteracted by increased output from the
ACL intact limb. Significant increases were observed in
the ACL intact limbs in all variables except for biceps
femoris and semitendinosis activation levels, which were
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Berchuck, M., Andriacchi, T.P., Bach, B.R., Reider, B., 1990. Gait
adaptations by patients who have a deficient anterior cruciate
ligament. J. Bone Joint. Surg. 72-A, 871–877.
Chmielewski, T.L., Rudolph, K.S., Fitzgerald, G.K., Axe, M.J.,
Snyder-Mackler, L., 2001. Biomechanical evidence supporting a
differential response to acute ACL injury. Clin. Biomech. 16, 586–
591.
Devita, P., Hortobagyi, T., Barrier, J., Torry, M., Glover, K.L.,
Speroni, D.L., Money, J., Mahar, M.T., 1997. Gait adaptations
before and after anterior cruciate ligament reconstruction surgery.
Med. Sci. Sports Exerc. 29, 853–859.
Ferber, R., Osternig, L.R., Woollacott, M.H., Wasielewski, N.J., Lee,
J., 2002. Gait mechanics in chronic ACL deficiency and subsequent
repair. Clin. Biomech. 17, 274–285.
Fleming, B.C., Beynnon, B.D., Renstrom, P.A., Peura, G.D., Nichols,
C.E., Johnson, R.J., 1998. The strain behaviour of the anterior
cruciate ligament during bicycling: an in vivo study. Am. J. Sports
Med. 26, 109–118.
Fleming, B.C., Renstrom, P.A., Beynnon, B.D., Engstrom, B., Peura,
G.D., Badger, G.J., Johnson, R.J., 2001. The effect of weightbearing and external loading on anterior cruciate ligament strain. J.
Biomech. 34, 163–170.
Gregor, R.J., Cavanagh, P.R., Lafortune, M., 1985. Knee flexor
moments during propulsion in cycling––A creative solution to
LombardÕs paradox. J. Biomech. 18, 307–316.
Imran, A., OÕConnor, J.J., 1998. Control of knee stability after ACL
injury or repair: interaction between hamstrings contraction and
tibial translation. Clin. Biomech. 13, 153–162.
Kvist, J., Gillquist, J., 2000. Anterior positioning of tibia during
motion after anterior cruciate ligament injury. Med. Sci. Sports
Exerc. 33, 1063–1072.
Lafortune, M.A., Cavanagh, P.A., Sommer, H.J., Kalenka, A., 1992.
Three-dimensional kinematics of the human knee during walking.
J. Biomech. 25, 347–357.
Limbird, T.J., Shiavi, R., Frazer, M., Borra, H., 1988. EMG profiles of
knee joint musculature during walking: changes induced by
anterior cruciate ligament deficiency. J. Orthop. Res. 6, 630–688.
Liu, W., Maitland, M.E., 2000. The effect of hamstring muscle
compensation for anterior laxity in the ACL-deficient knee during
gait. J. Biomech. 33, 871–879.
Marans, H.J., Jackson, R.W., Glossop, N.D., Young, C., 1989.
Anterior cruciate ligament insufficiency: a dynamic three-dimensional motion analysis. Am. J. Sports Med. 17, 325–332.
McLeod, W.D., Blackburn, T.A., 1980. Biomechanics of knee rehabilitation with cycling. Am. J. Sports Med. 8, 175–180.
McNair, P.J., Wood, G.A., Marshall, R.N., 1992. Stiffness of the
hamstring muscles and its relationship to function in the anterior
cruciate ligament deficient individuals. Clin. Biomech. 7, 131–137.
Mohr, T.M., Allison, J.D., Patterson, R., 1981. Electromyographic
analysis of the lower extremity during pedaling. J. Orthop. Sports
Phys. Ther. 2, 163–170.
More, R.C., Karras, B.T., Neiman, R., Fritschy, D., Woo, S.L.,
Daniel, D.M., 1993. Hamstrings––an anterior cruciate ligament
protagonist. Am. J. Sports Med. 21, 231–237.
Neptune, P.R., Kautz, S.A., 2000. Knee joint loading in forward
versus backward pedaling: implications for rehabilitation strategies. Clin. Biomech. 15, 528–535.
Nordeen-Snyder, K.S., 1977. The effect of bicycle seat height variations
upon oxygen consumption and lower limb kinematics. Med. Sci.
Sports Exerc. 9, 113–117.
Roberts, C.S., Rash, G.S., Honaker, J.T., Wachowiak, M.P., Shaw,
J.C., 1999. A deficient anterior cruciate ligament does not lead to
quadriceps avoidance gait. Gait Posture 10, 189–199.
Sanderson, D.J., Hennig, E.M., Black, A.H., 2000. The influence of
cadence and power output on force application and in-shoe
pressure distribution during cycling by competitive and recreational cyclists. J. Sports Sci. 18, 173–181.
van Lent, M.E.T., Drost, M.R., Wildenberg, F.A.J.M., 1994. EMG
profiles of ACL-deficient patients during walking: the influence of
mild fatigue. Int. J. Sports Med. 15, 508–514.
Winter, D.A., 1980. Overall principle of lower limb support during
stance phase of gait. J. Biomech. 13, 923–927.
Winter, D.A., 1983. Biomechanical motor patterns in normal walking.
J. Mot. Behav. 15, 302–330.
Winter, D.A., 1984. Pathologic gait diagnosis with computer-averaged
electromyographic profiles. Arch. Phys. Med. Rehabil. 65, 393–
398.
OF
406 References
407
408
409
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413
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