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The effect of an anterior cruciate ligament deficiency on steady-rate cycling biomechanics Hunt, Michael Anthony 2002

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THE EFFECT OF AN ANTERIOR CRUCIATE LIGAMENT DEFICIENCY ON STEADY-RATE CYCLING BIOMECHANICS  By  MICHAEL ANTHONY HUNT  B.H.K., The University of British Columbia, 2000  A THESIS SUBMITTED IN PARTIAL FULFILMENT OF THE REQUIREMENTS FOR THE DEGREE OF MASTER OF SCIENCE In THE FACULTY OF GRADUATE STUDIES (School of Human Kinetics) We accept this thesis as conforming to the required standard  THE UNIVERSITY OF BRITISH COLUMBIA  October 2002 © Michael Anthony Hunt, 2002  In presenting this thesis in partial fulfilment  of the  requirements for an advanced  degree at the University of British Columbia, I agree that the Library shall make it freely available for reference and study. I further agree that permission for extensive copying of this thesis for scholarly purposes may be, granted by the head of my department  or  by his  or  her  representatives.  It  is  understood  that  copying  or  publication of this thesis for financial gain shall not be allowed without my written permission.  Department of  IWwv^  KV\AAA^S>  The University of British Columbia Vancouver, Canada  DE-6 (2/88)  Abstract It is known that individuals missing a functional anterior cruciate ligament (ACL) in one limb exhibit changes in the walking biomechanics in that limb during mid-stance (10-30% of the gait cycle). Specifically, they exhibit reduced activation of the quadriceps muscle group and increased activation of the hamstring muscle group, resulting in a decreased net knee joint extensor moment and increased knee joint flexion. These compensations have been called a "quadriceps avoidance" strategy. It is not known whether these compensations are the direct result of the injury, or whether compensations made early in the rehabilitation process play a role in these changes. The purpose of the present study was to investigate the lower limb biomechanics of ACL deficient individuals during a common rehabilitation exercise for this injury - stationary cycling. Ten individuals with a unilateral ACL deficiency and ten age- and gendermatched controls performed six randomized bouts of stationary cycling for approximately one minute at intensities resulting from the combination of two cadences (60 and 90 rpm) and three power outputs (75, 125, and 175 W). It was found that, similar to during walking, ACL deficient individuals exhibited decreases in the magnitude of the quadriceps muscle activation in the injured limb. When combined with no change in hamstrings muscle activation, this resulted in a decreased net knee joint extensor moment in the injured limb. However, in contrast to walking, where increases from the hip or ankle extensors compensate for the decreased output from the knee joint extensors, ACL deficient individuals in the present study decreased output from the entire injured limb, resulting in a "limb avoidance". This limb avoidance was manifested by decreases in the magnitude of muscle activation from the rectus femoris,  n  vastus lateralis, and gluteus maximus, as well as decreases in the amount of force applied to the pedal. It was concluded that these compensations occurred in order to reduce anterior tibial translation in the injured limb. These results may suggest that a "quadriceps avoidance" strategy may be due in part to a "limb avoidance" strategy learned early during rehabilitation.  Table of Contents Abstract  ii  Table of Contents  iv  List of Tables  vi  List of Figures  vii  Acknowledgements  viii  Chapter 1: Introduction  1  Chapter 2: Methods  4  2.1 Subjects 2.2 Experimental Task 2.3 Instrumentation 2.4 Data Reduction 2.4.1 EMG Data 2.4.2 Pedal Force Data 2.4.3 Kinematics 2.4.4 Joint Moments of Force 2.5 Statistical Analysis  Chapter 3: Results 3.1 3.2 3.3 3.4  Linear Impulse Joint Moments of Force Electromyography Kinematics  Chapter 4: Discussion 4.1 4.2 4.3 4.4  Effects on the injured limb Effects on the uninjured limb Mechanisms Implications  4 5 5 7 7 8 9 9 10  12 12 13 16 20  21 21 25 27 29  Chapter 5: Conclusions  31  5.1 Future Directions  32  Chapter 6: References  34  IV  Appendix A: Literature Review A.l ACL Function and Biomechanics A. 1.1 ACL structure and function A. 1.2 ACL biomechanics A.2 Effects of an ACL Deficiency on Gait Biomechanics A.2.1 Lower Limb Kinetics A.2.2 Lower Limb Electromyographic Patterns A.2.3 Lower Limb Kinematic Patterns A.3 ACL Biomechanics and Cycling  Appendix B: Subject Data  38 38 38 38 40 40 43 44 47  49  Appendix B-l: Resultant pedal force Appendix B-2: Linear Impulse Appendix B-3: Joint Moments of Force Appendix B-3-1: Hip Joint Moments Appendix B-3-2: Knee Joint Moments Appendix B-3-3: Ankle Joint Moments Appendix B-4: EMG Ensemble Averages Appendix B-5: Integrated EMG (iEMG)  50 59 61 62 71 80 89 120  Appendix C: Informed Consent Form  126  v  List of Tables Table 2.1  Subject characteristics of the ACL group  4  Table 2.2  Subject characteristics of the control group  5  Table 2.3  Subject positioning during EMG normalizing trials. Hip angle is the included angle between the trunk and thigh, knee angle is the included angle between thigh and tibia, and ankle angle is the included angle between the tibia and the superior aspect of the foot. Strap placement for all trials was around the lower leg at the height of the lateral malleolus 7  Table 3.1  F- and p-values for limb effects as determined by repeated measures ANOVA  13  Group mean (SD) peak extensor and flexor moments for each of the six testing conditions. All values are in Nm. Note that a negative peak flexor moment at the ankle joint indicates the group average peak flexor moment was extensor in direction, indicates that ACL deficient limb peak moment was smaller than ACL intact  15  Group mean (SD) peak extension and flexion values (in degrees) for each of the three lower limb joints  20  Table 3.2  Table 3.3  vi  List of Figures Figure 2.1  Schematic diagram defining joint angles and joint moments. Arrows indicate direction of extensor moment about each joint... 10  Figure 3.1  Mean (SD) linear impulse of resultant pedal force for each of the six cycling intensities. Symbols from left to right for a given condition correspond to CON L (•), CON R (•), ACL I (A) and ACLD(o)  12  Rectus femoris iEMG data for all six conditions. Symbols from left to right for a given condition correspond to CON L (•), CON R (•), ACL I (A) and ACL D (o)  16  Vastus lateralis iEMG data for all six conditions. Symbols from left to right for a given condition correspond to CON L (•), CON R (•), ACL I (A) and ACL D (o)  17  Biceps femoris iEMG data for all six conditions. Symbols from left to right for a given condition correspond to CON L (•), CON R (•), ACL I (A) and ACL D (o)  17  Figure 3.2  Figure 3.3  Figure 3.4  Figure 3.5  Semitendinosis iEMG data for all six conditions. Symbols from left to right for a given condition correspond to CON L (•), CON R (•), ACL 1 (A) and ACL D (o)  18  Figure 3.6  Gluteus maximus iEMG data for all six conditions. Symbols from left to right for a given condition correspond to CON L (•), CON R (•), ACL I (A) and ACL D (o) 18  Figure 3.7  Ensemble average data for each of the five muscles at the 60 rpm /175 W cycling intensity. Lines indicate control (solid), ACL intact (dotted), and ACL deficient dashed)  19  Magnitude of anterior tibial translation during walking (in mm) with respect to % stride. Lines indicate control (solid), ACL intact (dashed) and ACL deficient (dotted)  23  Figure 4.1  Figure 4.2  Schematic representation of possible neural pathways associated with the reduction of quadriceps muscle activity in ACL deficient limbs as a result of increased anterior tibial translation. These pathways involve the interactions of la, lb (not shown), and II afferents synapsing with quadriceps alpha motor neuron inhibitory interneurons 28  vii  Acknowledgements I would like to thank many individuals who have helped and supported me during the past two years. First, I would like to thank my thesis supervisor, Dr. David Sanderson, whose knowledge and support were instrumental in the completion of this thesis. He provided the right mix of questions and answers which kept me focused and on the right path. I would also like to thank Dave for allowing me to be a part of the HKIN 363 team, which allowed me to find my interest for teaching. Thank you also to my other committee members, Dr. Timothy Inglis and Dr. Helene Moffet for agreeing to be part of this project and for their helpful guidance during some tricky times. Many thanks to my friends and family for their support during my university career to date. As well, thank you to my fellow graduate students for their help and support and our daily trips to the "SUB"! Last, but not least, I would like to thank Leslie who was there every step of the way, for the good and the bad, and for supporting me in my decisions.  viii  Chapter 1: Introduction The anterior cruciate ligament (ACL) provides a restraint to tibial movement by preventing excessive anterior tibial translation with respect to the femur. The ACL is the most commonly injured ligament in the body, and has been reported to occur in 30 out of every 100 000 individuals annually (Bollen, 1998). After an ACL rupture, the knee becomes less stable due to the increased movement of the bones comprising the joint. Compensations are therefore necessary in order to maintain near-normal stability of the joint, thereby preventing further injury. The term "quadriceps avoidance" was coined by Berchuck et al. (1990) based on the exhibition of a net knee joint flexor moment during the mid-stance phase of walking (10-30% of the gait cycle) in ACL deficient limbs, in contrast to a net knee joint extensor moment typically observed in healthy limbs. Although later studies failed to report a net flexor moment, some reported significant reductions in the magnitude of the knee joint extensor moment in ACL deficient limbs (Devita et al., 1997; Chmielewski et al, 2001; Lewek et al., 2002). Although EMG data were not collected in these studies, the notion of a "quadriceps avoidance" strategy was based on a reduction in the magnitude of activation of the quadriceps muscle group. Berchuck et al. (1990) believed that ACL deficient subjects preferentially reduced the activation of these muscles. For a given amount of hamstrings activation, this reduction in quadriceps activity would result in a net knee joint moment being more flexor in magnitude. Support for the reduction in quadriceps muscle activation in ACL deficient limbs during walking has come from later studies examining EMG characteristics in these individuals. Limbird et al. (1998) and van Lent et al. (1994) have reported that muscle  1  activation in the rectus femoris (Limbird), vastus lateralis (both) and vastus medialis (van Lent) was significantly reduced in ACL deficient limbs compared to uninjured individuals. Further, there have been reports of reductions in the amount of knee joint extension in ACL deficient limbs during mid-stance (Devita et al., 1997; Chmielewski et al., 2001). Therefore, although the term "quadriceps avoidance" was based solely on joint kinetic data, there is sufficient evidence that ACL deficient individuals do preferentially reduce the amount of quadriceps activation in the injured limb during walking. It has been suggested that these changes occur in order to limit the amount of increased anterior tibial translation prevalent in the absence of a functional ACL (Berchuck et al, 1990; Devita et al., 1997; Ferber et al., 2002). Since the quadriceps muscles insert via the patellar tendon at the tibial tuberosity, a component of the force from these muscles will act to translate the tibia anteriorly. As a result, quadriceps activation will decrease knee joint stability by increasing the amount of anterior tibial translation. Quadriceps activation is most detrimental at mid-stance during walking, since anterior tibial translation is already largest at this time in both healthy (Lafortune et al., 1992) and ACL deficient individuals (Marans et al., 1989; Kvist and Gillquist, 2000). Therefore, an effective strategy for maintaining knee joint stability in ACL deficient limbs during mid-stance would be to reduce the amount of quadriceps muscle activation. One issue regarding the compensations made by ACL deficient individuals during walking is whether the rehabilitation process plays a role in these changes. Stationary cycling is an integral part of all ACL rehabilitation programs since it increases knee joint range of motion, as well as muscular strength and endurance, yet it is not known how the lower limb musculature adapts in response to an ACL deficiency during the movement.  2  Stationary cycling is similar biomechanically to walking. In both movements, the primary goal - forward propulsion - requires output from both limbs, in alternating phases of propulsion and recovery. This is achieved by muscular output from the lower limb muscles (eg. quadriceps, hamstrings, gluteals) that are each active at similar times during the given cycle. The range of motion about each of the three lower limb joints are similar for both movements, and the joint moments of force patterns exhibit alternating phases of flexor and extensor. Therefore, it would seem possible that similar motor programs are responsible for the execution of the two movements. If one of the goals of a successful rehabilitation program is to modify existing motor programs to adapt to the injury, then it is possible that the changes observed in lower limb biomechanics in ACL deficient individuals during walking are the result of modifications made to its motor program during the rehabilitation process. The purpose of the present study was to examine the effects of an ACL deficiency on lower limb biomechanics during a common rehabilitation exercise - stationary cycling. Due to the similarities between the movements, it was thought that similar changes as those observed during walking would be observed during cycling. Specifically, it was hypothesized that A C L deficient limbs would exhibit decreased rectus femoris and vastus lateralis activity, increased biceps femoris, semitendinosis, and gluteus maximus activity, decreased knee joint extensor moments, and increased hip joint  extensor moments. The constraints placed upon the rider by the mechanics of the bicycle itself decrease the degrees of freedom possible in terms of lower limb kinematics. As a result, it was hypothesized that no differences would be observed in the lower limb kinematics between limbs.  3  Chapter?: Methods 2.1 Subjects Two groups of subjects were recruited to participate in the study. Ten individuals (5 male, 5 female) with a unilateral rupture of the anterior cruciate ligament comprised the A C L group (Table 2.1). These subjects were between the ages of 20 and 35, had experienced the injury no less than one month prior to testing, had no previous lower limb surgery on either limb, and were able to walk pain-free without the aid of crutches or a knee brace. Ten age- and gender-matched individuals with no history of knee injury or lower limb surgery comprised the control group (Table 2.2). A l l participants were briefed on the purpose and procedures of the experiment, and signed an informed consent form approved by the University of British Columbia Ethical Review Board (Appendix C) outlining experimental details. Table 2.1: Subject characteristics of the ACL group. Subject  Age  Sex  1 2 3 4 5 6 7 8 9 10  34 26 20 21 25 20 28 24 31 22  F M F F M M M F M F  Avg (SD)  25.1 (4.7)  Mass (kg) 59.7 67.6 54.0 64.5 84.1 72.7 63.6 67.3 87.3 59.1 69.0 (10.9)  Leg Length (cm) 86 88 85 83 94 103 93 89 90 85 90.1 (6.0)  Limb  Time (months)  Right Right Right Left Left Right Left Left Right Right 4 Left; 6 Right  4 4 8 9 3 8 15 49 2 5 10.7 (14.0)  4  Table 2.2: Subject characteristics of the control group. Subject  1 2 3 4 5 6 7 8 9 10 Avg(SD)  Age  Sex  22  F F F F M M M M M F  31 28 24 25 30 22 27 28 22  25.9 (3.4)  Mass (kg)  Leg Length (cm)  56.8  85  60.4  85  61.4  84  57.4  87  73.2  92  77.3  87  56.8  83  79.5  91  70.5  84  51.0  84  64.4 (9.9)  86.2 (3.1)  2.2 Experimental Task Subjects performed six randomized bouts of stationary cycling at intensities comprised of two cadences (60 rpm, 90 rpm) and three power outputs generally used in an ACL injury rehabilitation program (75W, 125W, and 175W). Subjects were instructed to take as long as needed to reach the desired cadence at the power output chosen. When subjects were able to maintain the correct cadence (+/- 5%), electromyographic (EMG), bi-directional pedal reaction force, and kinematic data were collected bilaterally for a total of 18 seconds, at which time the subject was told to stop pedaling. An adequate rest period, of at least 3 minutes, separated cycling bouts.  2.3 Instrumentation Subjects rode on a standard racing bicycle mounted on a Velodyne Trainer (Schwinn, Chicago, IL, USA) which enabled accurate 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  5  from the seat to the pedal at bottom-dead-center (BDC) was equal to the vertical distance from the subject's 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 (Therapeutics Unlimited, Iowa City, IA, USA), and 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). An isometric normalizing contraction was performed after electrode application to each muscle. Subjects' limbs were placed in a position so as to isolate the chosen muscle (Table 2.3), and required to provide a static 75N resistance as measured by a force transducer (Artech S-Beam, Riverside, CA, USA). The positions enabled each of the joints to be in a neutral position and in the middle of the joints' ranges of motion. 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  6  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 position of the pedal spindle. Raw coordinate data were acquired by digitizing each frame for the entire 18 second data collection period.  Table 2.3: Subject positioning during E M G normalizing trials. Hip angle is the included angle between the trunk and thigh, knee angle is the included angle between thigh and tibia, and ankle angle is the included angle between the tibia and superior aspect of the foot. Strap placement for all trials was around the lower leg at the height of the lateral malleolus. Muscle  Position  rectus femoris seated vastus lateralis seated biceps femoris seated semitendinosis seated gluteus maximus standing  Hip angle  90 90 90 90 180  Knee angle  90 90 90 90 180  Ankle angle  90 90 90 90 90  2.4 Data Reduction  2.4.1 EMG Data  EMG data from riding and the normalizing contractions were rectified and filtered using a 4 -order low-pass Butterworth filter with a cut-off frequency of 4 Hz. The th  average value from the middle 2 seconds of normalizing data was then obtained  7  (aNORM). A ratio of experimental E M G to aNORM was calculated for all data points of the 18 second data collection period. Top dead centre (TDC) pulses were then identified for each limb based on data from the analog TDC pulses and were used to partition data into pedal cycles (TDC to TDC). The propulsion phase was defined as the first 180 degrees of crank rotation (TDC to BDC), while recovery was defined as the last 180 degrees of crank rotation (BDC to TDC). Integrated E M G (iEMG) were then calculated for each muscle and each pedal cycle using all data points (ratios) within the cycle as defined by the TDC pulses. Subject i E M G averages and standard deviations were calculated based on i E M G data from 15 consecutive cycles. Data within each pedal cycle were also time-normalized to 360 data points (each time-normalized data point signifying 1 degree of crank revolution) and within-subject ensemble averages were calculated for each muscle as the average of 15 consecutive cycles at each of the 360 crank positions. Between-subjects limb averages and standard deviations of i E M G and ensemble averages were calculated based on the within-subject averages.  2.4.2 Pedal  Force  Data  A l l pedal force data (normal and shear) were filtered using a 4 -order low-pass th  Butterworth filter with a cut-off frequency of 4 Hz. Resultant forces were calculated as the resultant force vector of the filtered normal and shear forces. Top dead centre (TDC) pulses were then identified for each limb based on data from the analog TDC pulses and were used to partition data into pedal cycles (TDC to TDC). Linear impulseswere then calculated for each complete pedal revolution based on resultant force data. Within-  8  subject average linear impulse (and SD) was calculated based on the average linear impulse from 15 consecutive trials. Resultant force data within each pedal cycle were then time-normalized to 360 data points (each time-normalized data point signifying 1 degree of crank revolution) and within-subject ensemble averages were calculated as the average of 15 consecutive cycles at each of the 360 crank positions. Limb average and standard deviations of linear impulse and ensemble averages were calculated based on the within-subject averages.  2.4.3 Kinematics  Raw coordinate data were filtered using a 4 -order low-pass Butterworth filter th  with a cut-off frequency of 4 Hz and scaled to known real units. Joint angles were calculated for the hip, knee, and ankle and are defined in Figure 2.1. Analog TDC data, collected at 600 Hz, were then transformed to 60 Hz by taking every 10 data point. 60 th  Hz TDC data were then used to partition joint angle data into pedal cycles. Once pedal cycles were identified, joint angle data were normalized to 36 data points (each point signifying 10 degrees of crank revolution) and an ensemble average of 15 consecutive pedal cycles was calculated at each of the 36 data points. Peak joint angles in extension at all joints were then identified.  2.4.4 Joint Moments of Force  Scaled coordinate data were used to calculate segmental centers of mass (COM) coordinates, segmental linear and angular velocities and accelerations. These data were combined with pedal force data (once transformed to 60 Hz) and used to calculate net  9  joint moments of force at the hip, knee, and ankle using conventional inverse dynamics principles. Joint moment of force data were then portioned into pedal cycles based on 60 Hz TDC data. Moment data were then normalized to 36 data points and an ensemble average of 15 consecutive pedal cycles was calculated at each of the 36 positions. Peak joint moments were identified in both flexion and extension based on the ensemble averages. Between-subject peak moments averages were calculated based on withinsubject peaks.  Figure 2.1: Schematic diagram defining joint angles and joint moments. Arrows indicate direction of extensor moment about each joint  2.5 Statistical Analysis All variables were analyzed separately to identify differences between the four limb types (control left, control right, ACL intact, and ACL deficient) using a 4 (limb) x 2 (cadence) x 3 (power output) repeated measures ANOVA with repeated measures on  10  the last two factors. Significant F-ratios were further analyzed with a Tukey HSD hoc test and differences were classified based on a significance level of 0.05.  Chapter 3: Results 3.1 Linear Impulse Linear impulse was calculated as the time-based integral of resultant pedal force over a complete pedal cycle and then averaged across 15 consecutive cycles. A significant limb effect (difference between the four limb types) was observed (F3 36=l5.868, pO.OOl; Figure 3.1). Tukey post-hoc tests revealed that for all conditions, j  impulse in the A C L deficient limbs was significantly smaller than impulse in ACL intact limbs (p<0.01), while no differences existed in values between control limbs (p>0.882). 160  n  120  3> 80 3 Q.  E  40  75 W  125 W  60 rpm  175 W  75 W  125 W  175 W  90 rpm  Figure 3.1: Mean (SD) linear impulse of resultant pedal force for each of the six cycling intensities. Symbols from left to right for a given condition correspond to CON L (•), CON R (•), ACL I (A) and ACL D (o).  When examining the magnitude of the resultant force at a knee joint angle of 38 degrees (the mean joint angle found to be associated with the largest amount of ACL strain during cycling - see Fleming et al., 1998), significant differences existed between  12  limbs (F3 36 9.679, pO.OOl). Again, post-hoc analysis revealed that resultant force at that =  !  joint angle was larger in ACL intact limbs than ACL deficient limbs, while no differences existed between control limbs.  3.2 Joint Moments of Force Net joint moments of force were calculated for the hip, knee, and ankle joints. In general, ACL intact limbs exhibited the largest peak extensor moments at the hip, knee, and ankle as well as the largest peak flexor moments at the hip and knee. Conversely, ACL deficient limbs exhibited the smallest peak moments at all three joints. Although significant differences existed between ACL intact limbs and ACL deficient limbs, no differences were observed between control limbs. Limb effect values (ANOVA values based on differences between the four limbs) are summarized in Table 3.1, while actual group mean data are summarized in Table 3.2  Table 3.1: F - and p-values for limb effects as determined by a repeated measures A N O V A .  Joint  Direction  F-Value  p-value  Hip  Extensor  21.856  <0.001  Flexor  1.223  0.316  Extensor  1.882  0.150  Flexor  20.224  <0.001  Extensor  25.032  <0.001  Flexor  0.229  0.876  Knee  Ankle  Results from Table 3.1 indicate that limb differences existed at each joint and in only one direction. These directions were those that had the largest magnitude joint  13  moment for each joint during propulsion (TDC to BDC). On the other hand, no effects were observed in any of the peak moments predominant during recovery (BDC to TDC). Post-hoc analysis revealed that when significant differences existed, A C L intact and deficient limbs were always significantly different from each other. Further, differences existed between A C L intact limbs and control limbs (ankle and hip) as well as A C L deficient limbs and control limbs (all joints).  14  Table 3.2: Group mean (SD) peak extensor and flexor moments for each of the six testing conditions. All values are in Nm. Note that a negative peak flexor moment at the ankle joint indicates that the group average peak flexor moment was extensor in direction. * indicates ACL deficient limb peak moment was smaller than ACL intact. 75/60 125/60 175/60 75/90 125/90 175/90 Hip CONL CONR ACL I ACL D Knee CONL CONR ACL I ACL D Ankle CONL CONR ACL I ACL D  Extensor 114.4(20.0)  88.5 (13.9) 92.9(15.3) 106.3 (12.7) 76.1 (15.2)*  90.5 (12.9)*  141.8(18.1) 141.6(20.2) 154.9(9.1) 99.4 (9.3)*  8.8 (3.6)  7.9 (6.6)  8.8 (4.5)  7.4 (6.4)  11.1 (3.4) 7.6 (4.7)  116.4(19.9)  72.7 (14.7)  108.5 (16.1)  138.6(16.0) 139.5 (16.5) 153.3 (16.3) 109.1 (15.5)*  83.7(14.2)  115.5 (14.1)  105.7(15.5)  129.3 (17.5)  70.3 (10.6)*  84.1 (23.6)*  12.5 (8.0)  9.3 (3.2)  13.8(5.9)  17.2 (5.5)  16.4 (8.0)  10.6 (3.4)  14.9(6.1)  20.3 (6.7)  13.4 (5.6)  13.6(6.9)  15.3 (5.8)  17.9(7.1)  17.6(6.3)  8.5 (6.3)  9.2 (8.6)  11.4 (3.2)  12.1 (5.1)  13.7(6.0)  23.9(3.3)  33.6(4.9)  40.3 (4.0)  24.2 (3.2)  31.7(3.5)  40.1 (5.7)  23.6(3.3)  32.0 (4.0)  39.2 (3.3)  24.5 (4.2)  30.7(1.7)  38.3 (5.6)  28.2 (3.1)  37.3 (4.3)  45.1 (7.6)  32.6 (5.5)  35.3 (3.9)  41.4 (8.8)  19.4 (3.8)*  25.5 (3.9)*  30.6 (7.1)*  23.8 (4.4)*  25.1 (3.8)*  30.7 (5.7)*  135.3 (20.2)  Flexor Hip CONL CONR ACL I ACL D Knee CONL CONR ACL I ACL D Ankle CONL CONR ACL I ACL D  15.2 (8.6)  17.3 (11.0)  27.1 (15.3)  16.1 (9.3)  24.5 (16.1)  27.3 (19.1)  18.1 (11.0)  20.4(10.3)  32.1 (17.6)  27.1 (11.2)  25.7 (14.1)  35.7(12.7)  21.1 (6.8)  23.2 (6.6)  31.2(15.3)  27.5 (9.9)  25.5 (12.0)  29.8(11.1)  12.6 (9.6)  20.6(13.9)  31.7(11.1)  20.0 (6.6)  24.9 (6.6)  20.5 (9.3)  42.7 (9.6)  49.9 (9.0)  59.3 (8.1)  31.8(5.4)  42.2 (5.9)  59.4(11.1)  42.6 (9.2)  49.8 (6.2)  58.4 (7.6)  34.4 (6.0)  43.6 (6.0)  60.1 (9.7)  44.4(12.6)  57.4(11.8)  56.9(13.2)  43.6(11.9)  51.6(9.4)  64.8 (9.0)  26.3 (6.4)*  38.9(9.0)*  33.8(10.2)*  27.3 (6.5)*  32.8(8.1)*  39.5 (7.4)*  -2.1 (0.9)  -0.9 (0.7)  0.8(1.2)  -1.3 (0.8)  -1.7 (0.9)  -1.1(1.0)  -2.0(1.4)  -1.2(1.2)  0.3 (2.2)  -0.1 (1.4)  -1.5 (2.6)  0.1 (1.0)  -2.8 (2.1)  -1.0(2.0)  0.8 (2.7)  -1.3 (2.0)  -1.9(2.2)  -0.8 (2.3)  -2.5 (1.3)  -1.5(2.4)  0.4 (2.7)  -0.2(1.4)  -0.9(1.5)  -0.1 (2.1)  15  3.3 Electromyography Integrated EMG (iEMG) was calculated for each of the five muscles and is summarized in Figures 3.2 - 3.6. Significant differences existed between limbs for rectus femoris (F =21.867, p<0.001), vastus lateralis (F =14.529, pO.OOl), and gluteus 3;36  3;36  maximus (F 36=5.650, p=0.003). In all cases where significant differences existed, ACL 3j  deficient limbs exhibited significantly less activity than ACL intact as well as control limbs. ACL intact limbs exhibited significantly greater activity in the rectus femoris and vastus lateralis muscles. Between-subjects ensemble average data for a single condition is shown in Figure 3.7. It can be seen that although the magnitude of muscle activation differs between limbs, a similar temporal pattern exists.  1.5  S UJ T5 N re E  1  o  z  0.5  75 W  125 W  60 rpm  175W  75 W  125 W  175 W  90 rpm  Figure 3.2: Rectus femoris iEMG data for all six conditions. Symbols from left to right for a given condition correspond to CON L (•), CON R (•), ACL I (A) and ACL D (o).  16  2  1.5 O  S  UJ TS 0)  N  re E L_  o  z 0.5  75 W  125 W  175 W  75 W  60 rpm  125 W  175 W  90 rpm  Figure 3.3: Vastus lateralis iEMG data for all six conditions. Symbols from left to right for a given condition correspond to C O N L (•), C O N R (•), A C L I (A) and A C L D (o).  2  i  1.5  N  re E  1  o  z  0.5  75 W  125 W  60 r p m  175 W  75 W  125 W  175 W  90 rpm  Figure 3.4: Biceps femoris iEMG data for all six conditions. Symbols from left to right for a given condition correspond to C O N L (•), C O N R (•), A C L I (A) and A C L D (o).  17  1  s W  •a o 0.5 N E L_  o  z  75 W  125 W  175W  75 W  60 rpm  125 W  175 W  90 rpm  Figure 3.5: Semitendinosis iEMG data for all six conditions. Symbols from left to right for a given condition correspond to CON L (•), CON R (•), ACL I (A) and ACL D (o).  2 i  75 W  125 W  60 rpm  175 W  75 W  125 W  175W  90 rpm  Figure 3.6: Gluteus maximus iEMG data for all six conditions. Symbols from left to right for a given condition'correspond to CON L (•), CON R (•), ACL I (A) and ACL D (o).  18  GM  Crank Angle (Degrees)  Figure 3.7: Ensemble average data for each of the five muscles at the 60 rpm / 175 W cycling intensity. Lines indicate control (solid), A C L intact (dotted), and A C L deficient (dashed).  19  3.4 Kinematics Selected peak kinematic data is summarized i n Table 3.4. N o significant differences existed at any o f the three joints between limbs. Table 3.3: Group mean (SD) peak extension and flexion values (in degrees) for each of the three lower limb joints.  75/60  125/60  175/60  75/90  125/90  175/90  Extension Thigh CONL  111.7(2.7)  111.2 (2.6)  111.2 (2.6)  112.4 (2.7)  112.0(2.3)  111.6(2.0)  CONR  114.0(3.2)  113.2 (4.1)  113.3 (4.2)  114.6(3.9)  114.3 (3.0)  113.9(3.6)  I  115.1 (6.4)  113.6 (6.4)  113.8(5.7)  115.8(6.2)  115.3 (5.3)  115.4 (5.9)  ACL D  115.8(4.4)  113.7(4.1)  114.1 (3.7)  115.6 (3.8)  115.4 (3.8)  115.1 (3.2)  21.8(4.9) 21.1 (6.9)  20.0 (5.9) 19.0(8.2)  23.4 (5.7) 23.2 (8.2)  24.2 (9.7) 27.7 (5.2)  21.1 (10.5) 23.5 (7.0)  20.3 (5.3) 19.4 (8.4) 21.7(8.8)  22.1 (5.1) 21.6(7.7) 25.4 (8.2)  24.5 (5.1)  27.3 (8.9) 28.9(5.6)  23.1 (4.3) 22.8(5.7) 26.8 (7.2) 28.4 (5.4)  21.7(6.6) 22.4 (7.9) 20.3 (8.4) 18.1 (6.2)  21.6 (7.5)  21.7 (7.2)  22.6 (6.4)  22.4 (6.0)  22.5(6.1)  22.0 (9.4)  22.7 (9.5)  21.7(7.3)  21.9(6.2)  22.4 (7.3)  20.9(10.5)  20.1 (10.3)  19.1 (7.4)  17.9(7.4)  20.6 (7.6)  18.5 (9.1)  18.3 (10.3)  17.0(6.2)  17.6(6.0)  18.6(7.4)  ACL  Knee CONL CONR ACL I ACL D  27.2 (4.5)  Ankle CONL CONR ACL I ACL D  Flexion Thigh CONL  155.8(4.5)  155.0(3.4)  154.5 (3.6)  156.5 (4.2)  155.7(4.5)  156.0 (3.9)  CONR  158.7(2.5)  158.1 (3.5)  157.1 (3.6)  159.5 (3.3)  158.6(3.3)  158.8(3.1)  I  158.0(6.6)  156.5 (6.0)  155.3 (6.0)  158.4 (6.8)  158.1 (6.0)  158.0(3.9)  ACL D  158.5 (4.2)  157.1 (3.3)  155.9(3.8)  158.8(4.9)  158.7(5.3)  158.8(3.9)  CONL  100.2 (3.5)  99.5 (2.9)  99.3 (2.9)  100.5 (2.8)  100.0(3.4)  99.6(2.3)  CONR  ACL  Knee 100.5 (2.4)  99.6 (2.5)  99.4 (2.1)  100.8(2.6)  100.2 (2.2)  100.0 (2.3)  I  99.6 (5.0)  98.3 (5.2)  98.5(5.3)  100.2(4.6)  100.1 (4.2)  99.1 (4.7)  ACL D  100.9(5.0)  99.0 (4.7)  99.5 (5.3)  100.6 (4.5)  100.5 (5.3)  99.7 (4.0)  CONL  -6.9 (7.7)  -10.0(5.3)  -11.4 (4.5)  -6.5 (7.1)  -8.1 (6.8)  -6.2 (7.0)  CONR  -8.0(7.9)  -9.1 (8.1)  -12.3 (8.9)  -6.4 (9.8)  -8.3 (7.7)  -7.2 (8.7)  I  -5.0 (9.3)  -9.3 (11.0)  -14.7(7.4)  -4.7 (7.7)  -5.4 (8.3)  -6.1 (9.7)  ACL D  -6.3 (8.2)  -8.9 (8.1)  -12.7(6.6)  -5.4 (8.0)  -6.5 (7.8)  -3.9 (7.6)  ACL  Ankle  ACL  20  Chapter 4: Discussion The purpose of the present study was to examine whether a "quadriceps avoidance" strategy was employed by ACL deficient individuals during stationary cycling. Results from the present study indicated that ACL deficient individuals reduced quadriceps activation during cycling in the injured limb. When combined with no change in hamstrings muscle activity this resulted in decreases in the magnitude of the net knee joint extensor moment. However, they also reduced extensor output at the hip and ankle. As a result, a "limb avoidance" strategy was employed during cycling that was manifested by significant reductions in extensor output from the hip, knee, and ankle, resulting in decreased force transfer to the pedal (as shown in the impulse data).  4.1 Effects on the injured limb Decreases in extensor output from all three joints in ACL deficient limbs during cycling are in contrast to findings reported in gait literature, which show an increase in extensor output from either the hip or ankle to compensate for reduced knee extensor output (Berchuck et al., 1990; Chmielewski et al., 2001; Roberts et al., 1999). A limb avoidance strategy is not desirable during walking as lower limb stability would become compromised. A reduction in the knee joint extensor moment will decrease the magnitude of the lower limb support moment (Winter, 1980), resulting in decreased limb stability. As a result, ACL deficient subjects increased the magnitude of the hip joint or ankle joint extensor moment during walking, maintaining an adequate support moment. In cycling, however, a support moment is not necessary as the body is supported by the  21  bicycle seat. Therefore, a reduction in the magnitude of the knee joint extensor moment can exist in the absence of compensations from other joints. It has been suggested that the reduction in knee joint extensor output in ACL deficient limbs during walking occurs in order to reduce the amount of anterior tibial translation (Berchuck et al., 1990; Devita et al, 1997; Ferber et al, 2002). The principle function of the ACL is to prevent excessive anterior tibial translation with respect to the femur, ensuring knee joint stability. In the absence of a functional ACL, the tibia is allowed to move forward more freely during movement. As a result, the joint becomes less stable, and the probability of falling increases. Because of its insertion at the tibial tuberosity, the quadriceps muscle group, when active, will result in anterior tibial displacement. As a result, quadriceps activation in an ACL deficient limb is detrimental since it will act to increase the already large anterior tibial translation exhibited during mid-stance, resulting in further decreases in knee joint stability. Therefore, an effective strategy to maintain knee joint stability in ACL deficient limbs is to reduce quadriceps muscle activation at times when anterior tibial translation is known to be largest. It has been shown that anterior tibial translation is largest in ACL deficient limbs during walking at a knee joint angle of approximately 20 degrees of flexion (Kvist and Gillquist, 2001). It was also shown that the amount of anterior tibial translation between ACL deficient and ACL intact limbs was similar at all joint angles except those between 10 and 30 degrees (see Figure 4.1). It is this range of knee joint angles that are present during the mid-stance phase. Since the amount of anterior tibial translation is similar between injured and uninjured limbs except during this range of knee joint angles, and lower limb biomechanics are also similar at all knee joint angles except these, it seems  22  that a close relationship exists between anterior tibial translation and lower limb biomechanics in A C L deficient limbs.  0  25  50  75  100  Figure 4.1: Magnitude o f anterior tibial translation during walking (in mm)with respect to % stride. Lines indicate control (solid), A C L intact (dashed) and A C L deficient (dotted). (Source: K v i s t and Gillquist, 2001).  The amount of anterior tibial translation is not known during cycling, but can be estimated based on tibiofemoral shear force data (Neptune and Kautz, 2000) and A C L strain data (Fleming et al., 1998). As the tibia translates anteriorly, a shear force is created between the tibia and femur. The larger the translation, the larger the shear force. It has been shown that during cycling, tibiofemoral shear force is largest during propulsion (Neptune and Kautz, 2000). This is most likely due to the large pedal forces present as the rider pushes down on the pedal propagating up the limb since it has been shown that tibiofemoral shear force increases as the force at a foot-implement interface increases (Fleming et a l , 2001; Torzilli et al., 1994). A C L strain data would also suggest that anterior tibial translation during cycling is increased during propulsion. Fleming et al. (1998) reported that the A C L is strained during propulsion (and is largest at a knee joint angle of 38 degrees), but not during recovery. As the tibia moves forward in a healthy limb, the A C L stretches to limit this movement, creating strain within the ligament. The greater the anterior tibial translation, 23  the greater the strain within the ligament. Therefore, since ACL strain during cycling is largest during propulsion, it can be assumed that anterior tibial translation is also largest at that time. Linear impulse data from the present study indicated that ACL deficient subjects reduced the amount of force exerted to the pedal during propulsion - the period in which the forces were largest - and specifically, at the time when ACL strain (and hence, anterior tibial translation) was largest. This was interpreted as an attempt to reduce the amount of tibial translation within the limb. Since force production at the pedal is the end result of the activation of lower limb muscles, a decrease in force production must have occurred from a reduction in muscle output. This, in turn, reduced the magnitude of the joint moments of force. In the present study, these muscles were the knee extensors (rectus femoris and vastus lateralis) and the hip extensors (gluteus maximus). This would suggest that the reduction in the net knee joint extensor moment occurred as a result of decreased quadriceps activation, not increased hamstrings (flexor) activation. Increases in hamstrings activity have been reported during walking in ACL deficient limbs (Limbird et al., 1988; van Lent et al., 1994). Since the hamstrings muscles have been shown to be effective in limiting the amount of anterior tibial translation (McNair et al, 1992; More et al, 1993; Imran and O'Connor, 1998), an increase in hamstrings activation in ACL deficient limbs can be seen as an attempt to reduce tibial movement. However, no increase in hamstrings activity was observed in the present study. Therefore, ACL deficient individuals in the present study maintained tibial positioning by reducing quadriceps activation, not by increasing hamstrings activation.  24  No differences were observed in lower limb kinematics during the present study. Previous gait studies have reported increased knee joint flexion in ACL deficient limbs, presumably the result of a combination of decreased quadriceps and increased hamstrings activity (Devita et al., 1997; Roberts et al., 1999; Ferber et al., 2002). One of the fundamental differences between walking and cycling is that during cycling the path that the lower limb travels is dictated by such mechanical constraints as seat height, seat tube angle, and crank length, decreasing the number of kinematic degrees of freedom. Since these variables were held constant during the present study, the movement path was similar for all limbs, resulting in no significant differences in lower limb kinematics.  4.2 Effects on the uninjured limb ACL intact limbs in the present study exhibited significantly increased output compared to ACL deficient limbs. Specifically, there was increased muscle activity in the rectus femoris, vastus lateralis, and gluteus maximus, increased force exerted to the pedal, and increased magnitudes of the hip joint extensor, knee joint flexor, and ankle joint extensor moments of force. This is also in contrast to findings in some previous walking studies (Berchuck et al., 1990; van Lent et al., 1994). It has been reported that ACL intact limbs also reduce certain lower limb biomechanical characteristics during walking, van Lent et al. (1994) showed that quadriceps muscle activity was reduced in not only ACL deficient limbs, but the contralateral, ACL intact, limb as well. Berchuck et al. (1990) also reported decreased knee joint extensor moments in ACL intact limbs in addition to the presence of a knee joint flexor moment in ACL deficient limbs. These compensations are thought to occur in  25  an effort to maintain symmetry between the limbs. This results in a smoother, more efficient gait, and also ensures that the individual's gaze remains level during walking. This interlimb symmetry was, however, not possible in the present study. This was due to the intensity constraints placed upon the riders. Since riders were required to maintain a constant cadence at a given power output, the total amount of force applied to the chain must have remained constant. If ACL deficient limbs reduced the amount of force exerted to the pedal - and therefore, the chain as well - in an effort to limit the amount of anterior tibial translation, a concurrent increase in force production from the ACL intact limb was necessary. This resulted in a large asymmetry between limbs in most of the biomechanical characteristics examined. It was precisely this aspect of cycling mechanics that made a limb avoidance strategy possible. Although propulsion during walking is achieved by output from both limbs, the means by which each limb generates the force necessary for forward propulsion is essentially independent of each other. In contrast, the limbs of a cyclist are connected by the cranks. These cranks then connect to the chain, which ultimately propels the bicycle forward. The amount of force applied to the chain is, therefore, the summation of the forces applied by both limbs. This system allows for one limb to contribute the majority of the propulsive force, while allowing for a decrease in force application by the other limb.  26  4.3 Mechanisms Increased anterior tibial translation may result in the reduction of quadriceps muscle activity in A C L deficient limbs via at least three negative feedback loops (Figure 4.2) In the absence of a functional A C L , other structures must compensate for the reduction in tibial restraint. Due to its insertion at the posterior aspect of the tibia, the hamstrings muscle group is a prime candidate. Hamstrings activation has been shown to be effective in limiting the amount of anterior tibial translation (McNair et al., 1992; More et al., 1993; Imran and O'Connor, 1998). However, added stress is placed on the muscles as they must work harder to maintain tibial position. Liu and Maitland (2000) showed that the hamstrings in an A C L deficient limb must be activated to 50% of their maximum output in order to maintain the same tibial position as an A C L intact limb with no hamstrings activation. This increase in force production by the muscles will activate Golgi tendon organs residing within the muscle (Lundy-Ekman, 2002). This will increase the firing rate of lb afferents, which synapse in the spinal cord with lb inhibitory interneurons that can then inhibit information transfer via alpha motor neurons that supply the quadriceps muscle group. The end result is a decrease in quadriceps muscle activation. Anterior tibial translation can also result in the passive stretch of the hamstrings muscle group. This change in length of the muscle fibers would activate muscle spindles running in parallel with the fibers. Activation of muscle spindles could increase the firing rate of la afferents which can also synapse within the spinal cord with la inhibitory interneurons, ultimately decreasing the amount of quadriceps muscle activation.  27  Monosynaptic inhibition in this case may also act to increase the firing rates of alpha motor neurons supplying the hamstrings muscles, resulting in increased hamstrings muscle activation (Lundy-Ekman, 2002).  Figure 4.2: Schematic representation of possible neural pathways associated with the reduction of quadriceps muscle activity in ACL deficient limbs as a result of increased anterior tibial translation. These pathways involve the interactions of la, lb (not shown), and II afferents synapsing with quadriceps alpha motor neuron inhibitory interneurons. (source: Lundy-Ekman, 2002)  It has also been shown that joint receptors may play a role in the reduction of quadriceps muscle activity in ACL deficient limbs during walking. Torry et al. (2000) injected healthy knees with varying amounts of saline solution, resulting in knee joint distension. Subjects were then asked to walk at a freely chosen speed while lower limb biomechanics were examined. It was found that as knee joint distension increased, the magnitude of quadriceps muscle activation decreased. It was suggested that artificially distending the knee joint was similar to knee joint capsular distension in an ACL  28  deficient limb as a result of anterior tibial translation, and resulted in increased activation of knee joint receptors. These receptors may reside within the joint capsule itself, or on the surface, which respond to skin stretch. Activation of joint receptors typically signals the limits of a joint's range of motion. In which case, further muscle activation would be undesirable. As a result, joint receptor activation increases the firing rate of group II afferents which then synapse with inhibitory interneurons that inhibit alpha motor neuron activation, in this case motor neurons supplying the quadriceps muscles.  4.4 Implications The presence of a limb avoidance strategy in ACL deficient limbs during a cycling activity is important to recognize in many ways. First, stationary cycling is used in rehabilitation for ACL injuries as it is believed to combat the strength deficits typically associated with the injury. If patients exhibit a limb avoidance strategy during the movement in an effort to protect the integrity of the joint, then this exercise is not as successful as previously believed. Limiting the amount of force exerted to the pedals by the injured limb will just extend the length of time required for a full recovery from the injury. The large asymmetries created by the implication of a limb avoidance strategy might result in further injuries to other parts of the body. Interlimb strength asymmetries are known to cause injuries due to malalignment and compensating by the stronger limb. In a situation where the primary focus is to treat the injured knee, another injury is undesirable.  29  Another goal of rehabilitation is learning to cope with the injury. Clearly there are modifications to the existing motor program responsible for the execution of cycling in ACL deficient individuals. These are presumably learned early after the injury during rehabilitation and remain for at least four years after the injury (the range of injury times in the present study was 2 to 49 months). It is possible that if an ACL injured subject learns a limb avoidance strategy during rehabilitation, it could be manifested to a lesser extent during walking, resulting in the exhibition "quadriceps avoidance" strategy.  30  Chapter 5: Conclusions The present study is the first study to investigate lower limb biomechanics in ACL deficient subjects during stationary cycling. The purpose was to examine whether the "quadriceps avoidance" phenomenon - a strategy exhibited by ACL deficient individuals during walking, characterized by reduced quadriceps muscle activation and knee joint extensor moments - was also present in cycling. It was hypothesized that due to the similar muscle activation patterns exhibited in the lower limb musculature during both movements, that a similar compensation strategy - a "quadriceps avoidance" strategy - would also be observed during stationary cycling. The results indicated that ACL deficient subjects did exhibit reduced quadriceps activation and knee joint extensor moments. However, in contrast to walking, individuals with an ACL deficiency did not compensate for reduced knee joint extensor output by increasing output from either the hip or ankle. In fact, a "limb avoidance" strategy manifested by reduced quadriceps and gluteus maximus activity, decreased extensor moments at the hip and ankle, decreased flexor moments at the knee, and decreased linear impulse exerted at the pedal - was observed in the ACL deficient limbs. This was thought to occur for a combination of reasons. In order to reduce the increased anterior tibial translation that is present in an ACL deficient limb, a reduction in the amount of forward pull on the tibia can be achieved by decreasing quadriceps muscle activation. As a result, knee joint stability is maintained. Further, stresses placed upon structures at the posterior aspect of the tibia that must increase output to compensate for the loss of ACL function, such as the hamstrings muscle group, are reduced by limiting the amount of anterior tibial translation.  31  During walking, a sufficient support moment, comprised of the sum of the extensor moments about the hip, knee, and ankle, is required to maintain stability within the lower limb (Winter, 1980). Therefore, if extensor output from one joint is reduced, the extensor output at one of the other two joints must increase, or the lower limb stability becomes compromised. In cycling, however, the body is supported by the seat. As a result, the maintenance of a support moment is not necessary, and decreases in knee joint extensor output can occur without the need for compensations from the other two joints. Further, due to the mechanics of cycling, a reduction in the output from one limb can be accompanied by an increase in production from the contralateral limb. In the present study, this increase in force production from the uninjured limb in ACL subjects was necessary to counteract a decrease in force production by the ACL deficient limb in order to maintain the correct riding demands. The results of the present study provide evidence for a possible link between compensations made by ACL deficient individuals during rehabilitation and future compensatory strategies during walking. It is possible that by employing a "limb avoidance" strategy during cycling, the motor program responsible for walking is modified to encourage decreased knee joint extensor output.  5.1 Future Directions One of the limitations to this study was the cycling intensities used. Although the intensities were chosen to represent those used during a typical rehabilitation program, it is possible that a limb avoidance strategy was possible due to the relatively low cycling intensities. It is known that asymmetry - as measured by pedal force linear impulse -  32  increases as intensity decreases (Sanderson, 1990). This would suggest a reliance upon one limb, most likely the strongest. In the case of ACL deficient subjects, the injured limb is weaker than the intact one. As a result, a reliance upon the intact limb for force production is more likely at lower intensities. Therefore, cycling biomechanics in ACL deficient subjects should be examined at near maximum intensities, where the reliance upon one limb becomes more difficult. Although results from the present study may suggest that compensatory strategies employed during rehabilitation may play a role in later lower limb biomechanics, there is no direct evidence to support this claim. As a result, it would be beneficial to investigate lower limb biomechanics in cycling and walking concurrently during the early stages of rehabilitation and compare them with data collected from the same group of subjects at a later time. If a limb avoidance strategy used in cycling results in a quadriceps avoidance strategy in walking, changes will become more evident in walking as time increases postrehabilitation.  33  Chapter 6: References 1. Arms SW, Pope MH, Johnson RJ, Fischer RA, Arvidsson I, Eriksson E. The biomechanics of anterior cruciate ligament rehabilitation and reconstruction. American Journal of Sports Medicine. 12: 8-18, 1984. 2. Barrack RL, Skinner HB, Buckley SL. Proprioception in the anterior cruciate deficient knee. American Journal of Sports Medicine. 17: 1, 1989. 3. Baugher WH, Warren RF, Marshall JH, Joseph A. Quadriceps atrophy in the anterior cruciate insufficient knee. American Journal of Sports Medicine. 12: 192-195, 1984. 4. Berchuck M, Andriacchi TP, Bach BR, Reider B. Gait adaptations by patients who have a deficient anterior cruciate ligament. Journal of Bone and Joint Surgery. 72A: 871-877, 1990. 5. Birac RC, Andriacchi TP, Bach BR. Time related changes following ACL rupture. Transactions of the Orthopaedic Research Society. 1: 231, 1991. 6. Bollen S. Ligament injuries of the knee - limping forward? British Journal of Sports Medicine. 32: 82-84, 1998. 7. Carlsoo S, Nordstrand A. The coordination of the knee-muscles in some voluntary movements and in the gait in cases with and without knee joint injuries. Acta Chirurgica Scandinavica. 134:423-426, 1968. 8. Chmielewski TL, Rudolph KS, Fitzgerald GK, Axe MJ, Snyder-Mackler L. Biomechanical evidence supporting a differential response to acute ACL injury. Clinical Biomechanics. 16: 586-591,2001. 9. Devita P, Hortobagyi T, Barrier J, Torry M, Glover KL, Speroni DL, Money J, Mahar MT. Gait adaptations before and after anterior cruciate ligament reconstruction surgery. Medicine and Science in Sports and Exercise. 29: 853-859, 1997. 10. Durselen L, Claes L, Kiefer H. The influence of muscle forces and external loads on cruciate ligament strain. American Journal of Sports Medicine. 23: 129-136, 1995. 11. Ferber R, Osternig LR, Woollacott MH, Wasielewski NJ, Lee J. Gait mechanics in chronic ACL deficiency and subsequent repair. Clinical Biomechanics 17: 274-285, 2002. 12. Fleming BC, Beynnon BD, Renstrom PA, Peura GD, Nichols CE, Johnson RJ. The strain behavior of the anterior cruciate ligament during bicycling. American Journal of Sports Medicine. 26: 109-118, 1998.  34  13. Fleming BC, Renstrom PA, Beynnon BD, Engstrom B, Peura GD, Badger GJ, Johnson RJ. The effect of weightbearing and external loading on anterior cruciate ligament strain. Journal of Biomechanics. 34: 163-170, 2001. 14. Hamar D. Isokinetic cycle ergometer in rehabilitation after knee injuries. Medicine and Science in Sports and Exercise. 28: SI80, 1996. 15. Hirokawa S, Solomonow M, Lu Y, Lou Z, D'Ambrosia R. Anterior-posterior and rotational displacement of the tibia elicited by quadriceps contraction. American Journal of Sports Medicine. 20: 299-306, 1992. 16. Hsieh Y, Draganich LF, Ho SH, Reider B. The effects of removal and reconstruction of the anterior cruciate ligament on patellofemoral kinematics. American Journal of Sports Medicine. 26: 201-209, 1998. 17. Imran A, O'Connor JJ. Control of knee stability after ACL injury or repair: interaction between hamstrings contraction and tibial translation. Clinical Biomechanics. 13: 153-162, 1998. 18. Kalund S, Sinkjaer T, Arendt-Nielsen L, Simonsen O. Altered timing of hamstring muscle action in anterior cruciate ligament deficient patients. American Journal of Sports Medicine. 18: 245-248, 1990. 19. Kvist J, Gillquist J. Anterior positioning of tibia during motion after anterior cruciate ligament injury. Medicine and Science in Sports and Exercise. 33: 1063-1072, 2001. 20. Lafortune MA, Cavanagh PA, Sommer HJ, Kalenka. Three-dikmensional kinematics of the human knee during walking. Journal of Biomechanics. 25: 347-357, 1992. 21. Lewek M, Rudolph K, Axe M, Snyder-Mackler L. The effect of insufficient quadriceps strength on gait after anterior cruciate ligament reconstruction. Clinical Biomechanics. 17: 56-63, 2002. 22. Limbird TJ, Shiavi R, Frazer M, Borra H. EMG profiles of knee joint musculature during walking: changes induced by anterior cruciate ligament deficiency. Journal of Orthopaedic Research. 6: 630-638, 1988. 23. Liu W, Maitland ME. The effect of hamstring muscle compensation for anterior laxity in the ACL-deficient knee during gait. Journal of Biomechanics. 33: 871-879, 2000. 24. Lundy-Ekman L. Neuroscience: fundamentals for rehabilitation. 2 edition. W.B. Saunders Co. nd  35  25. MacDonald PB, Hedden D, Pacin O, Sutherland K. Proprioception in anterior cruciate ligament-deficient and reconstructed knees. American Journal of Sports Medicine. 24: 774-778, 1996. 26. MacWilliams BA, Wilson DR, DesJardins JD, Romero J, Chao EYS. Hamstrings cocontraction reduces internal rotation, anterior translation, and anterior cruciate ligament load in weight-bearing flexion. Journal of Bone and Joint Surgery. 17: 817822, 1999. 27. Marans HJ, Jackson RW, Glossop ND, Young MC. Anterior cruciate ligament insufficiency: a dynamic three-dimensional motion analysis. American Journal of Sports Medicine. 17: 325-332, 1989. 28. McLeod WD, Blackburn TA. Biomechanics of knee rehabilitation with cycling. American Journal of Sports Medicine. 8: 175-180, 1980. 29. McNair PJ, Wood GA, Marshall RN. Stiffness of the hamstring muscles and its relationship to function in anterior cruciate ligament individuals. Clinical Biomechanics. 17: 274-285, 1992. 30. More RC, Karras BT, Neiman R, Fritschy D, Woo SL, Daniel DM. Hamstrings-an anterior cruciate ligament protagonist. American Journal of Sports Medicine. 21: 231237,1993. 31. Neptune PR, Kautz SA. Knee joint loading in forward versus backward pedaling: implications for rehabilitation strategies. Clinical Biomechanics. 15: 528-535, 2000. 32. Nordeen-Snyder KS. The effect of bicycle seat height variations upon oxygen consumption and lower limb kinematics. Medicine and Science in Sports and Exercise. 9: 113-117, 1977. 33. Noyes FR, Barber SD. Simon R. High tibial osteotomy and ligament reconstruction in varus angulated, ACL-deficient knees. A two- to seven-year follow-up study. American Journal of Sports Medicine. 21: 2-12, 1993. 34. Pandy MG, Shelburne KB. Dependence of cruciate-ligament loading on muscle forces and external load. Journal of Biomechanics. 30: 1015-1024, 1997. 35. Pandy MG, Shelburne KB. Theoretical analysis of ligament and extensor-mechanism function in the ACL-deficient knee. Clinical Biomechanics. 13: 98-111, 1998. 36. Patel RR, Hurwitz DE, Andriacchi TP, Bush-Joseph CC, Back BR. Mechanisms by which patients with anterior cruciate ligament deficiency generate the "quadriceps avoidance gait". Presented at the 21 Annual Meeting of the American Society of Biomechanics. Clemson University, South Carolina, 1997. st  36  37. Renstrom P, Arms SW, Stanwyck TS, Johnson RJ, Pope MH. Strain within the anterior cruciate ligament during hamstring and quadriceps activity. American Journal of Sports Medicine. 14: 83-87, 1986. 38. Roberts CS, Rash GS, Honaker JT, Wachowiak MP, Shaw JC. A deficient anterior cruciate ligament does not lead to quadriceps avoidance gait. Gait and Posture. 10: 189-199, 1999. 39. Sanderson DJ. The influence of cadence and power output on asymmetry of force application during steady-rate cycling. Journal of Human Movement Studies. 19:1-9, 1990. 40. Sanderson DJ, Hennig EM, Black AH. The influence of cadence and power output on force application and in-shoe pressure distribution during cycling by competitive and recreational cyclists. Journal of Sports Sciences. 18: 173-181, 2000. 41. Shiavi R, Zhang L, Limbird T, Edmondstone M. Pattern analysis of electromyographic linear envelopes exhibited by subjects with uninjured and injured knees during free and fast speed walking. Journal of Orthopaedic Research. 10: 226236, 1992. 42. Torry MR, Decker MJ, Viola RW, O'Connor DD, Steadman JR. Intra-articular knee joint effusion induces quadriceps avoidance gait patterns. Clinical Biomechanics. 15: 147-159, 2000. 43. Torzilli PA, Deng X, Warren RF. The effect of joint compressive load and quadriceps muscle force on knee motion in the intact and anterior cruciate ligament sectioned knee. American Journal of Sports Medicine. 22: 105-112, 1994. 44. van Lent MET, Drost MR, Wildenberg FAJM. EMG profiles of ACL-deficient patients during walking: the influence of mild fatigue. International Journal of Sports Medicine. 15: 508-514, 1994. 45. Winter DA. Overall principle of lower limb support during stance phase of gait. Journal of Biomechanics. 13: 923-927, 1980.  37  Appendix A: Literature Review A . l A C L Function and Biomechanics  A. 1.1 ACL structure andfunction The anterior cruciate ligament (ACL) originates centrally upon the tibial plateau and runs superiorly and laterally to insert into the medial aspect of the lateral epicondyle of the femur. Because of its orientation within the knee joint, the ligament's principal function is to prevent excessive anterior movement of the tibia with respect to the femur and provides 85-95% of the total anterior tibial restraint (Pandy and Shelburne, 1998). Excision of the ligament has been shown to significantly increase the amount of anterior translation (Hsieh et al., 1998). Recent evidence has pointed to a sensory role for the ligament. It has been found that individuals with a ruptured ACL exhibit decreased joint position sense in the injured limb compared to the healthy limb (Barrack et al., 1988; MacDonald et al., 1996). This would suggest that the ACL increases the stability of the knee joint not only by preventing excessive movement of the bones comprising the joint, but by providing feedback about the position of the joint in space.  A. 1.2 ACL biomechanics As the tibia moves forward during movement, the ACL stretches to limit the range of this movement. This creates strain within the ligament and has been measured on cadavers using implanted strain gauges and simulated muscle loads. Under simulated concurrent quadriceps and hamstrings activation, it has been found that the strain within the ligament is largest at knee joint angles between 10 and 35 degrees of flexion (Arms et  38  al., 1984; Renstrom et al., 1986). Durselen et al. (1995) also found that under simulated quadriceps activation the ACL strain was largest at a knee flexion angle of 30 degrees (where 0 degrees is full extension). This means that within this range of knee joint angles, the resistance to anterior tibial displacement is largest. In the absence of a functional ACL, the lack of resistance will result in increases in the amount of tibial translation at times when strain in the functional ACL is largest. It has been shown in cadavers that differences in anterior tibial translation between ACL intact and ACL deficient knees occur only at knee flexion angles between 10 and 50 degrees, with the largest difference at approximately 30 degrees (More et al., 1993). It has also been shown that ACL strain - and hence, anterior tibial translation - is affected by the magnitude of the loading across the knee joint. Fleming et al. (2001) reported significant increases in ACL strain when subjects' knees were flexed to 20 degrees and placed in a device capable of producing a compressive load across the knee joint compared to no compressive force at all. Torzilli et al. (1994) used cadavers and compressive loads across the knee to show that the compressive knee load significantly increased anterior tibial translation compared to no compression at all knee joint angles. When the knee is flexed and a compressive force is applied across the knee joint, this force will not only result in a vertical compressive force, but a horizontal shear force acting to displace the tibia anteriorly. Therefore, the amount of ACL strain and subsequent anterior tibial translation in an ACL deficient limb is not a fixed value. It is dependent upon such factors as knee joint angle and the magnitude of the shear force across the knee joint. As a result, large  39  anterior tibial translation in an ACL deficient knee would be expected at times when the knee is loaded and flexed between 10 and 30 degrees (see Figure 4.1).  A.2 Effects of an A C L Deficiency on Gait Biomechanics It has been shown that a rupture of the ACL results in significant changes in the lower limb biomechanics during the mid-stance phase (10-30% of the cycle) of gait. Studies have investigated the lower limb kinetic, electromyographic, and kinematic patterns exhibited by ACL deficient individuals during this movement.  A.2.1 Lower Limb Kinetics The term "quadriceps avoidance" was introduced by Berchuck et al. (1990) to describe the knee joint moment profiles of ACL deficient individuals - 75% of whom exhibited a net flexor moment during midstance in their ACL deficient limb. This is in contrast to healthy individuals who show a net extensor moment. Although EMG data were not collected by Berchuck et al., it was postulated that the ACL injured individuals preferentially decreased the activity of the quadriceps femoris muscle group in the ACL deficient limbs, resulting in a net flexor moment. More recent studies investigating knee joint moment patterns in ACL deficient individuals found mixed results. Noyes et al. (1993) found a quadriceps avoidance pattern in 50% of subjects with an ACL deficiency. Birac et al. (1991) reported quadriceps avoidance in 80% of subjects tested greater than six years post-injury and 45% of subjects who were tested less than 1.5 years post-injury. Other studies have not found a net knee joint flexor moment, but did find a significantly reduced net knee joint extensor moment during midstance. Devita et al. (1997) examined biomechanical characteristics of  40  ACL deficient individuals before ligament reconstruction surgery and reported a decrease in the magnitude of the maximum knee joint extensor moment during midstance. Lewek et al. (2002) also found that ACL deficient subjects exhibit significantly reduced knee extensor moments during midstance. Chmielewski et al. (2001) reported that ACL deficient individuals who reported episodes of the injured knee giving-way (termed "noncopers") exhibited a significantly reduced maximum knee joint extensor moment during midstance. Other studies have reported no significant differences in the knee joint moment profiles between ACL deficient limbs and uninjured limbs (Roberts et al., 1999). Therefore, although a consensus cannot be reached on the exact changes exhibited by ACL deficient individuals in knee joint kinetics during walking, it is generally accepted that compensations are made by the knee joint extensor mechanism during the midstance phase of gait. One important aspect of walking is that during single limb stance, the weight of the body must be supported. This is generally accomplished by maintaining a rigid lower limb during stance and can be quantified by calculating a support moment which is the summation of the joint moments about the three lower limb joints (Winter, 1980). The sum of these moments must be extensor in direction in order to keep the individual upright. In the case of an ACL deficient individual who exhibits a reduced (or absent) knee joint extensor moment, one of two results may occur. First, all else being equal, the magnitude of the support moment will be decreased, resulting in reduced stability, and an increased probability of falling. This is clearly not desirable, especially in a clinical population such as ACL injured patients. The other option is to increase the magnitude of  41  the extensor joint moment about both or either of the other two joints. This will result in the maintenance an adequate support moment, and decrease the likelihood of falling. It has been reported that ACL deficient individuals counteract a decrease in the magnitude of the knee joint extensor moment by increasing the extensor moment at the hip joint (Berchuck et al, 1990; Patel et al., 1997). Devita et al. (1997) reported no differences in the maximum magnitude of the hip extensor moment, but found that the duration of the hip moment being extensor in direction was increased, resulting in a small increase in angular impulse about the hip joint. Chmielewski (2001) found that some ACL deficient subjects exhibited an increase in the magnitude of the ankle extensor moment. Compensations made by the ACL deficient limbs in response to the injury will result in a kinematic, kinetic, and electromyographic asymmetry between the limbs unless changes are made within the ACL intact limb as well. A functional symmetry is desirable to enhance the smoothness and efficiency of the movement. Although few studies have investigated lower limb biomechanics in ACL deficient individuals during walking bilaterally, a reduction in the magnitude of the knee joint extensor moment in the ACL intact limb has been reported (Berchuck et al. 1990). The notion of a "quadriceps avoidance" strategy was based on joint kinetic data and the muscular changes that would result in the differences in knee joint moment profiles (especially by the quadriceps muscle group) were assumed. As a result, later studies investigated electromyographical changes in ACL deficient individuals in order to get a better understanding of contributions of the knee joint musculature to the compensations made by ACL deficient individuals during walking.  42  A.l.2 Lower Limb Electromyographic Patterns The joint moment patterns exhibited in the lower limbs are the result of the interaction of the various muscles that cross the joints. For example, the knee joint moment profile is mainly the result of the interaction between the quadriceps muscle group - which extends the knee - and the hamstrings muscle group - which flexes the knee. A net knee joint extensor moment would therefore occur when the output of the quadriceps muscle group is larger than the hamstrings group. The ratio of these outputs will determine the magnitude of the joint moment. Since it has been shown that significant changes to the knee joint moment profile occur in ACL deficient limbs, it would seem reasonable to assume that changes in the EMG profile of the quadriceps and/or hamstrings muscles occur in these limbs as well. Although many studies have examined the muscle activation patterns of lower limb muscles in ACL deficient limbs (Carlsoo and Nordstrand, 1968; Kalund et al, 1990; Shiavi et al., 1992), few have measured muscle activation patterns in ACL deficient limbs and uninjured limbs concurrently during walking. Limbird et al. (1988) examined lower limb muscle activation patterns during a freely chosen walking speed and found significant decreases in the magnitude of rectus femoris and vastus lateralis activation and increases in biceps femoris and semitendinosis in ACL deficient limbs compared to uninjured limbs during midstance. Similar results were shown by van Lent et al. (1994) who reported significant decreases in vastus medialis and vastus lateralis in ACL deficient limbs compared to uninjured limbs. However, muscular activation was also measured in the ACL intact limbs within the injured group, and was reported to have  43  similar levels as the ACL deficient group, suggesting a compensation towards maintaining symmetry between limbs. Therefore, the reduced knee joint extensor moment exhibited in ACL deficient limbs occurs as a result of the combination of decreased quadriceps muscle activity and increased hamstrings activation. Further, a compensation at the hip joint resulting in a greater hip extensor moment can be explained by a greater tendency to extend the hip caused by increased hamstrings activation - and a lower tendency to flex the hip - caused by decreased quadriceps activation. Since the tendencies to flex and extend the hip and knee are altered due to changes in quadriceps and hamstrings muscle activity, changes in lower limb kinematics have also been reported in ACL deficient limbs during walking.  A. 2.3 Lower Limb Kinematic Patterns Increases in the amount of knee joint flexion have been reported in ACL deficient limbs compared to uninjured limbs during mid-stance (Devita et al., 1997; Roberts et al., 1999; Ferber et al, 2002). This can be perceived as a reluctance to extend the knee through the joint angle range which results in increased anterior tibial translation and can be caused by a reduction in the magnitude of quadriceps muscle activation. Increases in the amount of hip extension have also been reported in ACL deficient limbs (Berchuck et al., 1990; Ferber et al., 2002). An increase in the amount of hip extension can be caused by increased hamstrings muscle activation and may occur in conjunction with an increase in the magnitude of the net hip joint extensor moment reported in ACL deficient limbs (Berchuck et al., 1990; Patel et al, 1997; Devita et al., 1997).  44  A.l.4 Lower Limb Biomechanics and ACL Biomechanics It has been suggested that the reductions in quadriceps muscle activation and subsequent reductions in net knee joint extensor moments and knee extension observed in ACL deficient limbs during mid-stance are the result of increases in anterior tibial translation in the injured limb (Berchuck et al., 1990; Devita et al., 1997; Ferber et al., 2002). These changes occur at a time when the knee is flexed through a range of joint angles resulting in the largest anterior tibial translation in the ACL deficient limb. Also, large compressive forces are present within the knee during mid-stance as the single limb accepts the weight of the body and is the sole means of support. It is known that the largest amount of anterior tibial translation during walking in healthy individuals occurs during midstance (Lafortune et al., 1992). Other studies have shown that the largest difference in anterior tibial translation between ACL deficient subjects and uninjured subjects occurs during midstance (Kvist and Gillquist, 2000; Marans et al., 1989).Although it seems that an intimate relationship exists between lower limb biomechanics in ACL deficient limbs and anterior tibial translation, why are these changes manifested? In a healthy knee as the tibia moves forward normally, the ACL stretches to prevent excessive movement. However, in an ACL deficient limb, this restraint is absent, and the tibia is allowed to move forward more freely, resulting in decreased stability of the knee joint. Since it is known that increased quadriceps activation results in increased anterior tibial translation (Hirokawa et al, 1992; Arms et al., 1984; Renstrom et al., 1986; Pandy and Shelburne, 1997), a reduction in the magnitude of quadriceps activation would therefore decrease anterior tibial translation and maintain joint stability.  45  Increased anterior tibial translation may also be prevented by an increase in the amount of the rearward pull from structures at the posterior aspect of the tibia. Numerous studies have provided evidence suggesting that the hamstrings muscle group plays a protagonistic role to the ACL. ACL strain has been shown to decrease across all knee joint angles as the magnitude of the hamstrings activation is increased (Arms et al., 1984; Durselen et al., 1995). Also, studies investigating the magnitude of tibial movement in ACL deficient limbs have shown decreased anterior tibial translation as hamstrings activation is increased (More et al., 1993; Imran and O'Connor, 1998; Mac Williams et al, 1999). Clearly, changes in the amount of efferent drive to certain muscles are required to maintain tibial position in ACL deficient limbs during movement. The muscles mainly affected are the quadriceps and hamstrings muscle groups. Although these muscles are antagonistic muscles, they are connected by numerous feedback loops which control their activation levels during movement. Torry et al. (2000) have reproduced quadriceps avoidance-like characteristics (decreased quadriceps activation, decreased knee extension) by increasing intra-articular knee joint effusion. It was suggested that the knee joint capsular distention present was similar to that during periods of excessive anterior tibial translation and that negative feedback loops between knee joint receptors and alpha motor neurons supplying the quadriceps muscle group were responsible for the decrease in quadriceps activation. It might also involve feedback loops resulting in facilitation of hamstrings muscle activation to reduce the anterior tibial translation.  46  Liu and Maitland (2000) used a mathematical model to calculate hamstrings activation in an ACL deficient limb. It was reported that in order to maintain the amount of anterior tibial translation, the hamstrings muscles in an ACL deficient limb must produce 50% more force than muscles in an ACL intact limb. This increased force production, coupled with passive stretch from the anterior tibial displacement, would activate Golgi organs and muscle spindles within the hamstrings muscles and may result in reciprocal inhibition between the agonistic hamstrings and antagonistic quadriceps muscle groups. The decreases in quadriceps activation and increases in hamstrings activation needed to reduce anterior tibial translation are consistent with the findings of EMG studies of ACL deficient individuals during walking (Limbird et al., 1988; van Lent et al., 1994). This decrease in the ratio of quadriceps output to hamstrings output would also result in the increased knee flexion and a reduced knee joint extensor moment characteristic of a quadriceps avoidance gait.  A.3 A C L Biomechanics and Cycling Stationary cycling is an integral part of all rehabilitation programs for an ACL rupture both pre- and post-surgery. It enables the patient to increase knee range of motion, muscular strength and endurance, and cardiovascular endurance, while decreasing the potentially harmful ground reaction forces found in other activities (McLeod and Blackburn, 1980). To date, only one study has been reported that investigated any facet of cycling biomechanics in ACL injured subjects. Hamar (1996) reported substantial decreases in  47  the maximum effective force exerted in a 10 second maximum cycle at 60 rpm in ACL reconstructed knees (212 N) compared to the uninjured knees (423 N) following ACL reconstruction surgery. Following 5 weeks of rehabilitation, the maximum force increased to 760 N (reconstructed limbs) and 792 N (uninjured limbs). The asymmetry between limbs was likely due to the decrease in maximum strength commonly reported in ACL injured limbs (Baugher et al., 1984). However, it might suggest that differences may exist in the pattern of force application - and subsequent lower limb biomechanics between ACL injured and non-injured limbs during stationary cycling. Although the amount of anterior tibial translation is not known during cycling, it can be determined indirectly through ACL strain data, since the ACL strain profile and anterior tibial translation profile have been shown to be highly correlated (Pandy and Shelburne, 1998). Fleming et al. (1998) reported that ACL strain was largest during the propulsive phase of cycling, and smallest during recovery. This increased strain is likely due to the relatively large forces needed during propulsion resulting in increased tibial shear force (Neptune and Kautz, 2000). Therefore, if compensations are needed to reduce anterior tibial translation during cycling, they would most likely be observed during propulsion. A detailed analysis of lower limb biomechanics in ACL deficient individuals during cycling would provide valuable information for researchers and physical therapists. If the biomechanical changes associated with a quadriceps avoidance gait are present during cycling, then a new model for investigating this phenomenon is provided. Also, since stationary cycling is present in all rehabilitation programs for ACL injuries, the efficacy of this exercise can be determined.  48  Appendix B: Subject Data Individual subject data is included for the following variables: •  resultant pedal force  •  linear impulse of the resultant pedal force  •  joint moments of force (hip, knee, and ankle)  •  rectified, filtered EMG (rectus femoris, vastus lateralis, biceps femoris, semitendinosis)  •  integrated EMG (iEMG)  Abbreviations: •  CON = control subject  •  ACL = ACL injured subject  Appendix B-l: Resultant pedal force Resultant pedal forces were calculated as the resultant force vector comprised of the normal (Fz) and shear (Fy) forces applied to the pedal. Graph data denotes the ensemble average of 15 consecutive cycles (TDC to TDC) for each limb. Units on the vertical axis are N, while crank angle (in degrees) is on the horizontal axis.  Note: For all graphs, there will be two sets of data. For control subjects, the solid line indicates the left limb, while the dotted line indicates the right limb. For ACL subjects, the solid line indicates the intact limb, while the dotted line indicates the deficient limb.  50  u  u  u  u  o  o  o  o  o  a  a u  o  a o  IT)  r--  S a io O CO  -J  u  Q in  O o  ^  CO  Q «5  u<  O  Q Co T  O o co  U  Q ^  O  S  o  ^3"  CO  =8: -J  u  s  O  Q •>r  O co  4fc —1 U  53  S  o  §  ID  in  o  S T  o  CO  §  If)  s a  s-  o  00  U  -J  U  —1  U  OS  o =tfc  -J  54  u  u  u  u 55  56  IT)  c. s© 9\  5? T  CN  o CO  =tfc  U  u 57  in  E a  s©  0\ S  O  ^  ID  E a  s-  o ON  o ON  00  -J  U  ON  —1  <  u  3t —]  58  Appendix B-2: L i n e a r Impulse Linear impulse was calculated based on the integrated resultant force (Fr) for an entire cycle (TDC to TDC). Values given in the following tables are the average impulse and standard deviations over 15 consecutive cycles. Units on the vertical axis are Ns and crank angle (in degrees) on the horizontal axis.  Abbreviations: Control subjects: L = left limb R = right limb ACL injured subjects: I = intact limb D = deficient limb  59  Sub# 1  Limb L R 2 L R L 3 R 4 L R L 5 R L 6 R 7 L R 8 L R 9 L R 10 _ L R  60/75  60/125  60/175  90/75  90/125  90/175  98.0 (4.8)  112.9(4.8)  119.5 (14.4)  66.1 (3.1)  72.0 (5.3)  78.6 (3.3)  92.7 (5.1)  112.8(5.4)  110.5(14.0)  66.5 (2.6)  64.9 (4.2)  74.2 (3.5)  78.0 (6.5)  111.0 (7.4)  131.5 (6.0)  72.2 (3.9)  73.3 (6.7)  63.4 (6.6)  80.9 (6.1)  136.8(16.2)  129.2(19.4)  72.4 (7.2)  102.2 (7.2)  71.1 (8.1)  105.3 (10.5)  120.2(6.4)  145.9(7.4)  57.8(5.9)  74.6 (8.4)  100.1 (5.3)  122.3 (9.5)  143.4 (7.5)  179.4(12.9)  48.5 (6.6)  77.8 (2.7)  96.5 (3.7)  97.4 (4.8)  109.4 (9.0)  118.1 (8.9)  66.4 (4.8)  79.0 (6.5)  85.5 (6.6)  70.8(11.6)  100.6(6.8)  115.3 (9.5)  58.1 (9.6)  79.8(11.5)  80.8 (5.5)  72.0 (7.1)  88.6 (9.8)  140.3 (12.6)  70.2 (8.9)  97.0 (6.9)  104.3 (7.5)  99.0 (5.7)  99.8 (5.7)  149.6(11.6)  64.7 (9.0)  98.6(12.0)  110.1 (13.5)  113.5 (6.1)  140.7 (7.4)  125.2(10.8)  80.9 (7.8)  75.8 (7.2)  84.8 (9.7)  108.5 (9.1)  125.1 (6.6)  132.6(11.0)  86.4 (3.5)  77.3 (4.4)  85.7 (3.9)  69.2 (6.1)  85.6 (6.1)  110.9 (5.7)  52.5 (5.1)  88.8(8.5)  116.2(9.7)  57.1 (3.6)  69.9 (5.4)  125.5(6.5)  46.5 (3.2)  74.2 (4.7)  109.1 (9.1)  72.4 (6.5)  84.8 (4.6)  113.2(8.8)  57.4 (6.9)  80.3 (8.7)  110.9(11.7)  98.4 (7.3)  104.8(5.9)  102.7(4.3)  70.3 (3.9)  80.7 (4.6)  105.3 (10.9)  131.0(9.5)  158.1 (9.9)  138.9(12.0)  76.9 (6.8)  83.6(13.3)  108.5 (13.3)  101.1 (8.4)  111.8(5.5)  124.3 (12.2)  67.2 (5.3)  72.5 (5.4)  86.0(10.8)  103.5 (6.1)  134.0(8.3)  150.2(10.0)  62.7 (5.5)  99.1 (12.4)  116.8(9.4)  111.1 (10.1)  90.6 (9.7)  137.3(14.0)  49.3 (4.3)  87.7 (4.4)  90.1 (8.9)  Control Subjects Linear Impulse Sub# 1 2 3 4 5 6 7 8 9 10  Limb I D I D I D I D I D I D I D I D I D I D  60/75  60/125  60/175  90/75  90/125  90/175  111.9(11.4)  130.3 (6.2)  148.7(13.3)  68.5 (6.0)  72.3 (7.7)  75.3 (6.4)  75.3 (4.7)  70.5 (5.9)  80.9 (5.4)  41.1 (6.2)  62.6 (8.9)  58.2(10.8)  101.0(7.3)  115.0 (6.6)  144.4(11.6)  73.7 (2.8)  87.6 (2.9)  107.3 (2.0)  62.4 (3.7)  87.1 (6.0)  97.8 (6.5)  53.3 (2.0)  58.0 (2.9)  70.5 (4.0)  92.9(12.1)  147.4(5.7)  145.8(20.2)  76.2 (4.0)  91.9 (5.3)  90.8 (5.2)  60.0 (3.7)  83.8(7.1)  108.7(9.2)  53.7 (5.6)  64.6 (5.1)  89.0 (7.5)  115.7(3.5)  148.5 (8.1)  158.1 (17.4)  100.5(5.2)  94.5 (5.4)  94.5 (7.6)  87.5 (5.1)  122.3 (6.5)  115.6(13.8)  61.7(2.4)  69.0 (4.0)  71.3 (5.3)  135.2 (6.9)  149.7(11.9)  152.2(8.4)  99.0 (4.4)  106.0(8.7)  113.8(6.3)  82.2 (3.9)  86.2 (6.4)  94.6 (6.9)  69.5 (3.6)  65.9(4.1)  78.1 (5.4)  100.3 (9.3)  122.1 (11.8)  166.7(12.6)  82.5 (6.6)  100.0(9.7)  147.6(12.9)  68.0 (5.1)  106.0(5.0)  127.2(10.0)  53.7 (3.2)  76.9 (5.6)  84.5 (6.7)  139.3 (10.8)  155.1 (4.4)  141.7(19.2)  98.1 (5.4)  97.8 (4.9)  108.4 (7.8)  57.3 (4.3)  51.8(6.1)  50.2 (7.4)  70.5 (5.1)  83.6 (7.8)  83.1 (11.0)  120.9(11.5)  102.9(14.1)  148.2 (24.9)  87.7 (7.0)  98.3 (5.0)  101.8(5.0)  109.0 (6.2)  100.3 (6.6)  103.8(8.9)  46.6 (4.8)  43.9(6.6)  49.2 (4.3)  81.7(7.5)  111.3 (6.4)  121.7(9.1)  70.6 (7.8)  74.4 (5.1)  71.5 (7.1)  78.1 (6.3)  81.1 (7.1)  88.9(10.1)  56.9 (3.2)  52.9 (4.9)  66.6 (5.9)  81.9 (7.2)  137.1 (12.1)  149.1 (9.1)  59.2 (5.6)  99.9 (4.4)  99.5 (9.6)  89.8 (7.9)  115.4(16.7)  132.9(13.3)  66.5 (3.8)  86.2 (8.0)  110.5(14.1)  A C L injured Subjects Linear Impulse  60  Appendix B - 3 : Joint Moments of Force Joint moments of force were calculated were calculated for the hip, knee, and ankle joints. Graph data denotes the ensemble average of 15 consecutive cycles (TDC to TDC) for each limb. Directions for the extensor moments are given for each joint. Units on the vertical axis are Nm, while crank angle (in degrees) is on the horizontal axis.  Note: For all graphs, there will be two sets of data. For control subjects, the solid line indicates the left limb, while the dotted line indicates the right limb. For ACL subjects, the solid line indicates the intact limb, while the dotted line indicates the deficient limb.  61  Appendix B-3-1: Hip Joint Moments positive values = flexor moments negative values = extensor moments  in  a ©  in  ii  >  O  u  o  O  u  o  u  o  u  63  o  o  01  "1 j  in  E a  s-  o  r  J°  .co  < 0  {  0  /,*  CD / ,  9>  > 1  O U  o  u  00  ON  o  o  o  o  u  u  64  66  67  o  o  o  Q  O  9-  o  o  o  o  o  O  O  O  o O  o O  o  o  o  O  o  o  o  o  O  O  O  o  o  o  O  O  Q  o  o  o  o Q  o  o  o O  o Q  o  o  O  o O  o  o O  o  o  o Q  Q  o O  o  o O  o  O  o O  o  o  O  O  o Q  O  S a o ON  > II  so  o U  =tfc  o u  00  O  U  ON  =tfc  o u  o  z  o u  68  IT)  r--  © ON  IT)  ON  IT)  a © ON  ID  > -J  u  u 69  o O  o  o O  ^7  o  O (N  o Q  o  o  O  o  O CM  o  Q  o  o  O  o  O  V ^ N I  o  Q  o  o  Q ,  o  O CM  o  Q  o  o  o  V  O  O CN  70  Appendix B-3-2: Knee Joint Moments positive values = extensor moments negative values = flexor moments  o  \^  o ,7  1  ii  > +  in  O  u  o  o  o  4t  o 72  to  1 r  * / 5 a so  o  ° / '  •  CO  o • r-  '°f  IT)  a ©  4>  oo  ON  o  o  o  o  +  O  U  o  u  u  u  u  S-  o  s  >  +  °  s  g  s  s  o  3fc  g  s  °  a  =8:  u  u  -J  u  a  °  a  IT)  =tfc -J  u  74  S  9i >  +  O  Q i7>  O ^  Q ifi  O  Q uj)  O  ^  Q  in  O  Q  U*>  O  p  Q  iO  O  Q  oo  =tfc  U  U  Q  O l p  Q i O  O  Q  o  U*>  IT)  U  © ON  + X  I/)  o  o  u  o  u  o  u  o  u  76  77  S  ii >  +  °  S  g  8  0  S  g  S  0  S  |  S  0  S  g  S  °  S  |  3t  U  U 78  79  A p p e n d i x B-3-3: A n k l e Joint Moments positive values = plantarflexor moments negative values = dorsiflexor moments  in  a o  o t P u) 17  8  I  T> CN  up ^ r ^ - C M  &  o  is) CN  Co  to CN  Q Co  <S) r-p  m CM  D P^-  m CM  L  &  o  is) CN  ^  o CN  (~, Co  1  is) 1^-  t o CN  ^  m CN  0  Co  o  o CM  ^ o LO  CM  o LO  o  ^ I  ^ CN  o  ^ CN  o CT)  ^ t^-  LO CM  o  10 t\j  o J  S  *n  4t  o  ^  o  o  u  o  o  u  81  0 >  > X  so  O  u  O  u  oo  ON  o  O U  u  o 4t  O  u  82  84  o to  CO  o , CN  o  ,\  °  /'  at  /'  L  in  a o ON  If)  E a  u © ON  4t  o  o  u  o  u  o  u  o  u  85  in a © ON  >  =8:  O u  o  oo  ON  o  o  o  o  u  =tfc z  u  86  ^  CM  C  N  L  f  l  ^  C  N  CN  Co  I"-  CM  CN  Co  ^  CN  CN  to  87  88  Appendix B-4: E M G Ensemble Averages Ensemble averages were calculated for each subject and all five muscles. The averages are calculated based on 15 consecutive cycles (TDC to TDC). The values on the vertical axis are the ratio of task EMG (EMG data collected during the experimental task) to the average EMG required to maintain the 75 N isometric contraction for each muscle, while values on the horizontal axis are crank angles (in degrees). All EMG data (task and normalizing) were rectified and filtered using a 4 -order Butterworth filter with a lowth  pass cutoff frequency of 4 Hz. Each plot contains data for all subjects for the group. The top half contains data for the left / intact limb, while the bottom half contains data for the right / deficient limb for the control / ACL groups. Although values on the bottom half of the plots are negative, it is the absolute magnitude which is of importance.  Note: Data for G M was not obtained for A C L subject #2 due to technical difficulties.  Abbreviations  PvF = rectus femoris VL = vastus lateralis BF = biceps femoris ST = semitendinosis GM = gluteus maximus  89  CON R F 6 0 rpm/75 W Left Limb  6  CON RF 60 rpm/125 W Left Limb  91  92  CON RF 90 rpm/75 W Left Limb  Deficient Limb 180  270  360  CON RF 90 r p m / 1 2 5 W Left Limb  Deficient Limb 180  270  360  10  CON V L 6 0 rpm/75 W Left Limb  CON VL 60 r p m / 1 2 5 W  10  Left Limb  Deficient Limb  -10 0  90  180  270  CON VL 60 rpm/175 W Left Limb  Right Limb 180  270  360  270  360  ACL V L 6 0 rpm/175 W Intact Limb  Deficient Limb 180  10  CON VL 90 rpm/75 W Left Limb  99  10  CON VL 90 r p m / 1 2 5 W Left Limb  10  CON VL 90 rpm/175 W Left Limb  102  5  CON B F 6 0 rpm/125 W Left Limb  5  CON BF 60 rpm /175 W Left Limb  104  5  CON B F 9 0 r p m / 7 5 W Left Limb  106  CON ST 60 rpm/75 W Left Limb  270  360  270  360  A C L ST 60 rpm/ 75 W Intact Limb  Deficient Limb 90  180  108  4  CON ST 60 r p m / 1 2 5 W Left Limb  109  4  CON ST 60 rpm/175 W Left Limb  110  CON ST 90 rpm/75 W Left Limb  360  A C L ST 90 rpm/ 75 W Intact Limb  Deficient Limb 90  180  270  360  111  4  CON ST 90 r p m / 1 2 5 W Left Limb  112  4  CON ST 90 rpm /175 W Left Limb  113  CON GM60 rpm/75 W Left Limb  Right Limb 90  180  270  360  270  360  A C L GM 60 rpm/75 W Intact Limb  Deficient Limb 90  180  114  6  CON GM60 rpm/125 W Left Limb  115  116  6  CON GM90 rpm/75 W Left Limb  117  6  CON GM90 rpm/125 W Left Limb  118  119  Appendix B-5: Integrated E M G (iEMG) Integrated EMG (iEMG) was calculated for each subject for all five muscles. The integral is calculated based on the average of 15 consecutive time-based integrals for a complete cycle (TDC to TDC). The values used were the ratio of task EMG (EMG data collected during the experimental task) to the average EMG required to maintain the 75 N isometric contraction for each muscle. All EMG data (task and normalizing) were rectified and filtered using a 4 -order Butterworth filter with a low-pass cutoff frequency th  of 4 Hz before the integrals were calculated.  Note: Data for G M was not obtained for A C L subject #2 due to technical difficulties.  Abbreviations  RF = rectus femoris VL = vastus lateralis BF = biceps femoris ST = semitendinosis GM = gluteus maximus  120  Sub#  Limb  60/75  60/125  60/175  90/75  90/125  90/175  1  L  0.59 (0.09)  0.82 (0.12)  1.04 (0.18)  0.42 (0.07)  0.40 (0.09)  0.68 (0.12)  R  0.50 (0.08)  0.80 (0.10)  1.03 (0.20)  0.27 (0.03)  0.34 (0.07)  0.67 (0.11)  L  0.53 (0.04)  0.90(0.11)  1.17(0.12)  0.63 (0.12)  0.59 (0.12)  0.81 (0.13)  R  0.77 (0.13)  0.95 (0.18)  1.51 (0.13)  0.77 (0.08)  0.69 (0.09)  1.00 (0.16)  L  0.91 (0.19)  1.24 (0.18)  1.01 (0.19)  0.61 (0.11)  0.81 (0.14)  1.04 (0.23)  R  0.92 (0.09)  1.29 (0.11)  1.06 (0.12)  0.47 (0.07)  0.70 (0.05)  1.13 (0.13)  L  0.83 (0.10)  1.06 (0.12)  1.03 (0.07)  0.82 (0.11)  0.76 (0.09)  1.00 (0.12)  R  0.80 (0.11)  1.18(0.22)  1.31 (0.21)  0.79 (0.17)  0.66 (0.11)  1.10(0.17)  L  0.92 (0.17)  1.13 (0.17)  1.40 (0.18)  0.76 (0.14)  0.63 (0.09)  0.91 (0.19)  R  0.75 (0.10)  1.10(0.12)  1.34 (0.21)  0.74 (0.12)  0.79 (0.10)  0.96 (0.13)  L  1.05 (0.14)  1.03 (0.08)  1.36 (0.17)  0.83 (0.09)  1.03 (0.11)  0.93 (0.09)  R  1.18(0.17)  1.17(0.20)  1.43 (0.21)  0.82 (0.10)  0.94 (0.10)  1.06 (0.17)  L  0.84 (0.25)  1.14(0.20)  1.34 (0.22)  0.89 (0.14)  0.83 (0.10)  1.07 (0.17)  R  0.70 (0.17)  1.02 (0.22)  1.22 (0.20)  0.75 (0.08)  0.86 (0.12)  0.94 (0.12)  L  0.94 (0.18)  1.15(0.23)  1.51 (0.18)  0.61 (0.13)  0.83 (0.09)  1.10(0.13)  R  0.98 (0.12)  1.09 (0.10)  1.69 (0.12)  0.58 (0.09)  0.78 (0.16)  1.19(0.16)  L  1.31 (0.18)  1.21 (0.12)  1.47 (0.15)  0.90 (0.15)  0.77 (0.13)  0.95 (0.12)  R  1.17 (0.14)  1.35 (0.18)  1.42 (0.16)  1.01 (0.14)  0.67 (0.13)  1.03 (0.21)  L  0.72 (0.06)  1.08 (0.10)  1.60 (0.13)  0.60 (0.08)  0.79 (0.13)  1.04 (0.16)  R  0.71 (0.16)  1.15 (0.12)  1.51 (0.11)  0.61 (0.09)  0.82 (0.07)  1.11 (0.09)  2 3 4 5 6 7 8 9  10  Control Subjects RF i]EMG Sub#  Limb  60/75  60/125  60/175  90/75  1  90/125  90/175  I D I D I D I D I D I D I D I D I D I D  1.37(0.13)  1.72 (0.18)  2.16(0.19)  0.92 (0.15)  1.02 (0.20)  1.38(0.19)  0.69 (0.06)  0.79 (0.08)  1.14(0.08)  0.45 (0.06)  0.46 (0.06)  0.65 (0.08)  2 3 4 5 6 7 8 9  10  1.43 (0.26)  1.49 (0.23)  1.37(0.25)  0.79 (0.15)  0.87 (0.24)  1.00 (0.24)  0.93 (0.10)  0.82 (0.08)  0.91 (0.11)  0.62 (0.07)  0.52 (0.06)  0.68 (0.06)  1.14(0.23)  1.42(0.19)  1.74 (0.35)  0.71 (0.17)  0.87 (0.19)  0.92 (0.16)  0.47 (0.10)  0.67 (0.11)  0.92 (0.14)  0.41 (0.10)  0.43 (0.14)  0.79 (0.18)  1.24 (0.23)  1.49 (0.14)  1.90 (0.25)  1.11 (0.15)  1.04 (0.17)  1.48 (0.13)  0.70 (0.07)  0.94 (0.11)  1.29 (0.24)  0.53 (0.06)  0.65 (0.10)  1.00 (0.10)  1.38(0.18)  1.44 (0.20)  1.46 (0.18)  1.05 (0.09)  1.12(0.10)  1.19(0.12)  0.75 (0.10)  0.80 (0.07)  1.03 (0.08)  0.43 (0.05)  0.77 (0.11)  0.87 (0.15)  1.18(0.19)  1.50 (0.26)  1.66 (0.22)  0.95 (0.25)  1.14(0.28)  1.42 (0.23)  0.55 (0.09)  0.95 (0.12)  1.27 (0.20)  0.56 (0.15)  0.71 (0.12)  0.81 (0.14)  1.09 (0.22)  1.29 (0.27)  1.78 (0.17)  1.07 (0.18)  1.28 (0.25)  1.53 (0.44)  0.50 (0.11)  1.00 (0.35)  1.59 (0.33)  0.46 (0.07)  0.59 (0.06)  0.63 (0.12)  1.25 (0.20)  1.36 (0.20)  1.89 (0.28)  1.05 (0.17)  1.00 (0.15)  1.16(0.22)  0.78 (0.09)  0.88 (0.15)  1.33(0.17)  0.57(0.10)  0.62 (0.12)  0.67 (0.09)  0.86 (0.19)  1.24 (0.09)  1.34 (0.11)  0.63 (0.07)  0.64 (0.08)  0.72 (0.09) 0.77 (0.11)  0.70 (0.09)  0.80 (0.13)  0.94 (0.14)  0.64 (0.06)  0.72 (0.09)  0.88 (0.13)  1.34 (0.14)  1.58 (0.20)  0.92 (0.14)  0.89 (0.10)  1.05 (0.09)  0.62 (0.09)  0.93 (0.13)  1.28 (0.24)  0.71 (0.14)  0.69 (0.10)  0.81 (0.10)  A C L injured Subjects RF iEMG  121  Sub# 1 2 3 4 5 6 7 8 9 10  Limb L R L R L R L R L R L R L R L R L R L R  60/75  60/125  60/175  90/75  90/125  90/175  0.89 (0.08)  1.15(0.10)  1.21 (0.16)  0.71 (0.07)  0.67 (0.09)  0.89 (0.07)  0.82 (0.10)  1.11 (0.13)  1.26 (0.16)  0.68 (0.06)  0.68 (0.06)  0.91 (0.10)  1.04(0.10)  1.53 (0.14)  1.81 (0.13)  0.98 (0.08)  1.00 (0.09)  1.09 (0.11)  0.94(0.13)  1.14(0.17)  1.66 (0.13)  0.86 (0.09)  0.84 (0.12)  0.98 (0.10)  0.82 (0.10)  1.37(0.14)  1.52 (0.12)  0.98 (0.13)  0.93 (0.09)  1.06 (0.10)  0.65 (0.09)  1.48 (0.10)  1.70 (0.15)  0.85 (0.10)  0.86 (0.10)  0.99 (0.09)  0.75 (0.08)  0.96 (0.10)  1.17(0.11)  0.82 (0.11)  0.79 (0.10)  0.99 (0.13)  0.82 (0.11)  1.16(0.11)  1.21 (0.15)  0.73 (0.14)  0.66 (0.12)  0.92 (0.12)  0.91 (0.10)  1.19(0.15) 1.05 (0.12)  1.36 (0.15)  0.67 (0.13) 0.75 (0.14)  0.76 (0.08)  0.77 (0.09)  0.72 (0.09)  1.06 (0.16) 0.85 (0.09)  0.71 (0.12)  1.25 (0.13)  1.42 (0.17)  1.10(0.09)  1.22 (0.10)  1.52 (0.13)  1.04(0.13) 0.94 (0.10)  1.02 (0.12)  0.92 (0.10)  0.87 (0.06)  1.11 (0.10)  0.61 (0.08)  0.90 (0.08)  1.42 (0.17)  0.70 (0.11)  0.64 (0.07)  0.98 (0.15)  0.53 (0.07)  0.87 (0.08)  1.48 (0.16)  0.62 (0.08)  0.59 (0.07)  0.96 (0.11)  0.62 (0.11)  0.97 (0.11)  1.04 (0.12)  0.77 (0.15)  0.78 (0.10)  0.98 (0.15)  0.26 (0.04)  0.45 (0.06)  0.46 (0.05)  0.67 (0.13)  0.92 (0.11)  0.84 (0.14)  1.07 (0.20) 1.01 (0.12)  1.46(0.10) 1.35(0.16)  1.71 (0.15) 1.58 (0.12)  0.94 (0.08) 1.00 (0.11)  0.97 (0.11)  1.15 (0.08)  1.01 (0.11)  1.22 (0.08)  0.81 (0.12)  1.08 (0.08)  1.45 (0.12)  0.75 (0.09)  0.80 (0.10)  1.06 (0.09)  0.80 (0.19)  1.16(0.08)  1.61 (0.13)  0.90 (0.13)  0.82 (0.08)  0.97 (0.09)  1.23 (0.13)  Control Subjects V L iEMG Sub# 1 2 3 4 5 6 7 8 9 10  Limb I D I D I D I D I D I D I D I D I D I D  60/75  60/125  60/175  90/75  90/125  90/175  1.16(0.15)  1.21 (0.16)  1.69 (0.22)  1.17(0.26)  1.13 (0.21)  1.82(0.31)  0.80 (0.11)  0.82 (0.11)  1.04 (0.13)  0.64 (0.10)  0.67 (0.11)  0.84(0.12)  1.22 (0.18)  1.22 (0.08)  1.65 (0.23)  1.31 (0.10)  1.32 (0.14)  1.63 (0.12)  0.51 (0.05)  0.82 (0.07)  0.97 (0.07)  0.59 (0.04)  0.54 (0.04)  0.74 (0.08)  1.24 (0.41)  1.21 (0.13)  2.05 (0.31)  1.42 (0.26)  1.46 (0.25)  1.85 (0.18)  0.58 (0.16)  0.97 (0.12)  0.99 (0.13)  0.87 (0.14)  0.73 (0.15)  1.18(0.15)  0.89 (0.10)  1.29 (0.09)  1.81 (0.23)  0.98 (0.11)  0.87 (0.13)  1.24 (0.10)  0.53 (0.05)  0.83 (0.08)  1.21 (0.21)  0.53 (0.09)  0.47 (0.09)  0.75 (0.09)  1.16(0.18)  1.20 (0.78)  1.64 (0.16)  1.64 (0.16)  1.62 (0.22)  1.70 (0.21)  0.69 (0.06)  0.88 (0.09)  0.98 (0.07)  0.73 (0.05)  0.72 (0.06)  0.83 (0.06)  1.01 (0.17)  1.04 (0.18)  1.67 (0.18)  0.70 (0.11)  1.08 (0.14)  1.50 (0.17)  0.56 (0.09)  0.72 (0.10)  1.28 (0.19)  0.63 (0.13)  0.65 (0.12)  0.90 (0.12)  1.14(0.21)  1.34 (0.20)  1.44(0.17)  0.74 (0.09)  0.72 (0.08)  0.91 (0.10)  0.57 (0.08)  0.58 (0.05)  1.02 (0.12)  0.64 (0.11)  0.59 (0.06)  0.70 (0.11)  0.94 (0.20)  1.13 (0.23)  1.43 (0.15)  0.92 (0.14)  0.89 (0.13)  0.98 (0.12)  0.66 (0.12)  1.23 (0.21)  1.02 (0.14)  0.61 (0.13)  0.61 (0.14)  0.68 (0.09)  0.94 (0.16)  1.09 (0.09)  1.23 (0.10)  0.80 (0.15)  0.79 (0.08)  0.94 (0.10)  0.63 (0.13)  0.91 (0.09)  1.08 (0.15)  0.73 (0.07)  0.65 (0.09)  0.77 (0.10)  0.88 (0.12)  1.09 (0.14)  1.51 (0.14)  0.90 (0.14)  0.88 (0.07)  1.00 (0.13)  0.56 (0.09)  0.74 (0.10)  1.15(0.10)  0.64 (0.06)  0.62 (0.07)  0.77 (0.11)  A C L injured Subjects V iEMG  122  Sub# 1 2 3 4 5 6 7 8 9 10  Limb L R L R L R L R L R L R L R L R L R L R  60/75  60/125  60/175  90/75  90/125  90/175  0.45 (0.05)  0.73 (0.07)  1.00 (0.14)  0.43 (0.06)  0.47 (0.07)  0.73 (0.10)  0.33 (0.03)  0.54 (0.09)  0.73 (0.11)  0.33 (0.06)  0.39 (0.07)  0.46 (0.09)  0.47 (0.07)  0.83 (0.09)  1.14(0.11)  0.54 (0.06)  0.53 (0.09)  0.83 (0.11)  0.43 (0.09)  0.58 (0.07)  0.78 (0.11)  0.41 (0.05)  0.39 (0.06)  0.50 (0.06)  0.52 (0.11)  0.46 (0.04)  0.78 (0.13)  0.50 (0.06)  0.48 (0.07)  0.58 (0.06)  0.45 (0.07)  0.32 (0.06)  0.82 (0.11)  0.37 (0.09)  0.43 (0.05)  0.43 (0.12)  0.22 (0.04)  0.59 (0.07)  0.32 (0.06)  0.30 (0.06)  0.32 (0.08)  0.27 (0.03)  0.37 (0.06) 0.28 (0.04)  0.41 (0.03)  0.35 (0.06)  0.36 (0.07)  0.34 (0.06)  0.98 (0.09) 0.77 (0.09)  1.24 (0.09) 0.98 (0.08)  1.55 (0.13)  0.87 (0.11)  0.88 (0.08)  0.97 (0.22)  1.29 (0.20)  0.74 (0.11)  0.84 (0.07)  0.98 (0.13)  0.54 (0.08)  0.73 (0.10) 0.82(0.12)  1.32 (0.17)  0.54 (0.10)  0.58 (0.09)  0.84 (0.11)  0.61 (0.14)  1.14(0.22)  0.67 (0.15)  0.63 (0.09)  0.74 (0.16)  0.52 (0.16)  0.77 (0.18)  0.44 (0.13)  0.66 (0.30)  0.54 (0.12)  1.01 (0.23) 1.13 (0.32)  0.46 (0.25)  0.38(0.11)  0.31 (0.06)  0.34 (0.05)  0.35 (0.04)  0.71 (0.09) 0.63 (0.10)  0.74 (0.12)  1.20 (0.12)  0.34 (0.16)  0.90 (0.07)  0.36(0.10)  0.54 (0.13) 0.64 (0.09)  0.45 (0.07)  0.57 (0.08)  0.73 (0.15) 0.60 (0.21)  0.76 (0.15) 0.67 (0.22)  0.84 (0.15) 0.85 (0.21)  0.46 (0.17) 0.49 (0.15)  0.48 (0.12)  0.65 (0.14)  0.56 (0.09)  0.73 (0.24)  0.50 (0.10)  0.72 (0.11)  0.93 (0.14)  0.48 (0.08)  0.48 (0.13)  0.59 (0.06)  0.53 (0.09)  0.79(0.18)  1.12(0.11)  0.46 (0.08)  0.44 (0.04)  0.61 (0.11)  0.44 (0.06)  Control Subjects BF i]EMG Sub# 1 2 3 4 5 6 7 8 9 10  Limb I D I D I D I D I D I D I D I D I D I D  60/75  60/125  60/175  90/75  90/125  0.68 (0.12) 0.61 (0.04)  90/175  0.89 (0.09) 0.83 (0.09)  1.35 (0.08) 1.07 (0.10)  0.86 (0.08) 0.64 (0.07)  0.89 (0.15) 0.69 (0.06)  1.03 (0.13) 0.87 (0.12)  0.99 (0.08)  1.40 (0.21)  1.88 (0.13)  1.11 (0.10)  1.00 (0.07)  1.54 (0.14)  1.09 (0.16)  1.40 (0.20)  1.79 (0.13)  0.79 (0.06)  0.75 (0.04)  1.19(0.13)  0.23 (0.07)  0.40 (0.06)  0.54 (0.07)  0.23 (0.04)  0.20 (0.05)  0.35 (0.06)  0.36 (0.05)  0.46 (0.05)  0.58 (0.07)  0.26 (0.03)  0.25 (0.04)  0.27 (0.03)  0.80 (0.14)  0.88(0.10)  1.38(0.18)  0.54 (0.05)  0.55 (0.08)  0.72 (0.09)  0.33 (0.05)  0.52 (0.11)  0.79 (0.18)  0.33 (0.05)  0.32 (0.05)  0.40 (0.07)  0.61 (0.08)  0.71 (0.09)  1.02 (0.14)  0.53 (0.08)  0.34 (0.06)  0.62 (0.08)  0.61 (0.07)  0.64 (0.08)  0.87 (0.08)  0.55 (0.05)  0.63 (0.04)  0.63 (0.05)  0.72 (0.20)  1.45 (0.43)  1.29 (0.30)  0.72 (0.19)  0.85 (0.19)  0.94 (0.26)  0.52 (0.17)  0.61 (0.13)  0.92 (0.24)  0.63 (0.21)  0.73 (0.23)  1.00 (0.24)  0.67 (0.10)  1.18(0.09)  0.82 (0.11)  0.60 (0.07)  0.88 (0.17)  1.04 (0.13)  0.64 (0.13)  1.14(0.21)  2.27 (0.35)  0.62 (0.07)  0.95 (0.26)  1.21 (0.20)  0.54 (0.11)  0.64 (0.09)  0.82 (0.08)  0.55(0.13)  0.56 (0.14)  0.98 (0.12)  0.58(0.14)  0.64 (0.11)  0.75 (0.11)  1.09 (0.34)  0.86 (0.25)  1.08 (0.22)  0.57 (0.18)  0.73 (0.28)  1.16(0.28)  0.78 (0.08)  0.98 (0.17)  0.81 (0.15)  0.46 (0.08)  0.64 (0.19)  1.24 (0.14)  0.64 (0.13)  0.47 (0.06)  0.83 (0.22)  0.53 (0.13)  0.75 (0.13)  1.08 (0.12)  0.39 (0.08)  0.57 (0.14)  0.86 (0.15)  0.44 (0.09)  0.67 (0.08)  0.95 (0.16)  0.37(0.10)  0.56 (0.11)  0.74 (0.20)  A C L injured Subjects B iEMG  123  Sub# 1 2 3 4 5 6 7 8 9 10  Limb L R L R L R L R L R L R L R L R L R L R  60/75  60/125  60/175  90/75  90/125  90/175  0.30 (0.04)  0.36(0.03)  0.61 (0.08)  0.19(0.04)  0.21 (0.02)  0.29 (0.03)  0.29 (0.08)  0.43 (0.07)  0.54 (0.10)  0.24 (0.05)  0.28 (0.08)  0.34 (0.03)  0.30 (0.06)  0.44 (0.06)  0.59 (0.06)  0.37 (0.05)  0.32 (0.08)  0.59 (0.07)  0.35 (0.07)  0.49 (0.05)  0.65 (0.09)  0.44 (0.05)  0.43 (0.06)  0.56 (0.08)  0.39 (0.05)  0.37 (0.05)  0.55 (0.09)  0.25 (0.04)  0.34 (0.05)  0.34 (0.04) 0.32 (0.04)  0.41 (0.05)  0.36 (0.05)  0.51 (0.07)  0.26 (0.03)  0.32 (0.03)  0.25 (0.08)  0.36 (0.06)  0.46 (0.05)  0.36 (0.07)  0.32 (0.04)  0.30 (0.05)  0.34 (0.07)  0.54 (0.14)  0.88 (0.17)  0.39 (0.14)  0.42 (0.12)  0.45 (0.08)  0.23 (0.05) 0.22 (0.05)  0.29 (0.10) 0.24 (0.10)  0.52 (0.11) 0.47 (0.12)  0.13 (0.07) 0.13 (0.04)  0.28 (0.06) 0.29 (0.11)  0.26 (0.08)  0.17(0.04)  0.25 (0.05)  0.55 (0.09)  0.21 (0.09)  0.27 (0.10)  0.22 (0.06)  0.19(0.05)  0.30 (0.05)  0.48 (0.08)  0.23 (0.05)  0.26 (0.06)  0.29 (0.07)  0.30 (0.06)  0.34 (0.07)  0.62 (0.08)  0.24 (0.04)  0.27 (0.04)  0.51 (0.12)  0.38 (0.05)  0.41 (0.05)  0.82 (0.13)  0.30 (0.05)  .0.30 (0.04)  0.65 (0.06)  0.54 (0.07)  0.59 (0.07)  0.77 (0.07)  0.32 (0.11)  0.54 (0.08)  0.49 (0.09)  0.44 (0.09)  0.37 (0.09)  0.67 (0.08)  0.36 (0.07)  0.57 (0.07)  0.42 (0.08)  0.39 (0.06) 0.28 (0.12)  0.44 (0.09) 0.50 (0.12)  0.60 (0.08) 0.71 (0.10)  0.35 (0.06) 0.33 (0.10)  0.34 (0.08) 0.30 (0.09)  0.48 (0.09) 0.51 (0.08)  0.29 (0.05)  0.52 (0.08) 0.31 (0.04)  0.50 (0.08) 0.41 (0.04)  0.33 (0.06)  0.32 (0.04)  0.37 (0.04)  0.37 (0.05)  0.33 (0.03)  0.46 (0.05)  0.24 (0.03)  0.30 (0.10)  Control Subjects ST i E M G Sub# 1 2 3 4 5 6 7 8 9 10  Limb I D I D I D I D I D I D I D I D I D I D  60/75  60/125  60/175  90/75  90/125  0.29 (0.05)  90/175  0.37 (0.05)  0.58 (0.07)  0.39 (0.04)  0.55 (0.06)  0.20 (0.03)  0.31 (0.04)  0.44 (0.05)  0.30 (0.03)  0.40 (0.08) 0.32 (0.03)  0.21 (0.05) 0.24 (0.05)  0.42 (0.10) 0.46 (0.11)  0.80 (0.12)  0.24 (0.04)  0.16(0.02)  0.91 (0.12)  0.65 (0.08)  0.33 (0.05)  0.34 (0.06)  0.47 (0.09)  0.32 (0.08)  0.44 (0.06)  0.64 (0.08)  0.25 (0.05)  0.44 (0.08)  0.50 (0.07)  0.38 (0.06)  0.55 (0.08)  0.60 (0.05)  0.28 (0.04)  0.27 (0.05)  0.34 (0.04)  0.20 (0.03)  0.26 (0.06)  0.39 (0.05)  0.16(0.04)  0.17(0.03)  0.27 (0.03)  0.15 (0.05)  0.18(0.04)  0.31 (0.08)  0.12(0.02)  0.14(0.02)  0.14(0.02)  0.49 (0.08)  0.40 (0.08)  0.89 (0.16)  0.50 (0.08)  0.83 (0.12)  0.62 (0.10)  0.23 (0.07)  0.40 (0.16)  0.71 (0.15)  0.18(0.04)  0.31 (0.05)  0.22 (0.07)  0.39 (0.13)  0.35 (0.06)  0.23 (0.07)  0.24 (0.07)  0.24 (0.05)  0.35 (0.08) 0.27 (0.04)  0.23 (0.07)  0.22 (0.06)  0.22 (0.04)  0.21 (0.06)  0.22 (0.04)  0.23 (0.02)  0.35 (0.03)  0.54 (0.05)  0.20 (0.02)  0.27 (0.03)  0.31 (0.05)  0.36 (0.05)  0.51 (0.07)  0.77 (0.08)  0.34 (0.04)  0.48 (0.07)  0.52 (0.08)  0.43 (0.06)  0.42 (0.07)  0.51 (0.08)  0.56 (0.07)  0.29 (0.08)  0.28 (0.05)  0.29 (0.05)  0.21 (0.04)  0.26 (0.04)  0.29 (0.05)  0.19(0.03)  0.15(0.03)  0.19(0.02)  0.54 (0.11)  0.27 (0.04)  0.37 (0.08)  0.35 (0.04)  0.34 (0.05)  0.32 (0.06)  0.32 (0.04)  0.37 (0.06)  0.60 (0.07)  0.44 (0.09)  0.37 (0.06)  0.46 (0.08)  0.31 (0.04)  0.69 (0.12)  0.58 (0.08)  0.48 (0.11)  0.44 (0.11)  0.57 (0.12)  0.20 (0.04)  0.32 (0.05)  0.48 (0.08)  0.32 (0.09)  0.30 (0.05)  0.32 (0.05)  A C L injured Subjects ST i E M G  124  Sub# 1 2 3 4 5 6 7 8 9 10  Limb L R L R L R L R L R L R L R L R L R L R  60/75  60/125  60/175  90/75  90/125  90/175  0.50 (0.06)  0.89 (0.08)  1.47 (0.27)  0.50 (0.07)  0.59 (0.08)  1.19(0.17)  0.48 (0.04)  0.79 (0.08)  1.27 (0.26)  0.47 (0.04)  0.59 (0.07)  1.16(0.12)  0.62 (0.07)  1.03 (0.12)  1.28 (0.22)  0.52 (0.08)  0.65 (0.12)  0.53 (0.12)  0.95 (0.14)  1.24 (0.17)  0.48 (0.09) 0.64 (0.12)  0.40 (0.07)  0.74 (0.12)  0.39 (0.08)  0.55 (0.05)  0.74 (0.07)  0.52 (0.06)  0.40 (0.05)  0.46 (0.05)  0.46 (0.07)  0.66 (0.07)  0.94 (0.07)  0.45 (0.13)  0.46 (0.04)  0.53 (0.03)  0.77 (0.12)  0.97 (0.22)  1.51 (0.16)  0.74 (0.18)  0.73 (0.10)  0.96 (0.15)  0.64 (0.12)  0.84 (0.17)  1.24(0.24)  0.85 (0.17)  0.80 (0.21)  1.09 (0.24)  0.43 (0.05) 0.48 (0.07)  0.59 (0.07)  0.96 (0.14)  0.35 (0.04)  0.43 (0.06)  0.58 (0.11)  0.74 (0.16)  0.76 (0.14)  0.38 (0.08) 0.84 (0.12)  0.82 (0.08)  1.93 (0.22) 1.78 (0.12)  0.98 (0.08)  1.03 (0.14)  1.21 (0.15)  1.14(0.09)  1.48 (0.16) 1.58(0.17)  1.10(0.08)  1.01 (0.08)  1.14(0.13)  0.63 (0.14)  1.01 (0.09)  1.13 (0.14)  0.59 (0.06)  0.57 (0.09)  0.68 (0.06)  0.77 (0.08)  1.02 (0.08)  1.22 (0.04)  0.60 (0.05)  1.02 (0.08)  0.72 (0.02)  0.63 (0.09)  1.38 (0.36)  1.61 (0.22)  0.91 (0.21)  1.18(0.23)  1.27 (0.32)  0.62 (0.10)  1.15 (0.19)  1.37 (0.16)  0.98 (0.15)  0.88 (0.09)  0.99 (0.20)  0.46 (0.14) 0.52 (0.10)  1.08 (0.13) 1.05 (0.16)  1.38(0.28)  0.49 (0.07)  0.54 (0.12)  0.73 (0.12)  1.43 (0.23)  0.50 (0.08)  0.49 (0.09)  0.72 (0.10)  0.45 (0.09) 0.42 (0.12)  0.80 (0.10) 0.76 (0.08)  1.17(0.14)  0.44 (0.09) 0.40 (0.04)  0.46 (0.09)  0.59 (0.09)  0.48 (0.06)  0.60 (0.10)  1.35 (0.16)  0.90 (0.15)  Control Subjects G M iEMG Sub# 1 2 3 4 5 6 7 8 9 10  Limb I D I D I D I D I D I D I D I D I D I D  60/75  60/125  60/175  90/75  90/125  90/175  2.36 (0.38)  3.59 (0.32)  3.71 (0.47)  2.58 (0.13)  2.37(0.31)  3.23 (0.30)  1.25 (0.16)  . 2.39 (0.36)  2.58 (0.33)  1.47 (0.23)  1.46 (0.23)  2.12(0.03)  0.53 (0.15)  0.94 (0.12)  1.43 (0.26)  0.33 (0.09)  0.76 (0.10)  1.26(0.18)  0.61 (0.05) 0.56 (0.11)  0.59 (0.11) 0.54 (0.09)  0.95 (0.10) 0.82 (0.10)  .  0.74 (0.10)  1.07 (0.10)  1.37 (0.23)  0.93 (0.10)  0.79 (0.12)  0.94 (0.09)  0.57 (0.09)  0.76 (0.11)  0.96 (0.18)  0.71 (0.10)  0.58 (0.06)  0.76 (0.10)  0.91 (0.14)  0.77 (0.17)  1.03 (0.11)  0.43 (0.10)  0.54 (0.15)  0.88(0.15)  0.79 (0.09)  0.74 (0.10)  0.67 (0.06)  0.39 (0.07)  0.56 (0.11)  0.40 (0.05)  0.83 (0.13)  0.93 (0.16)  1.14(0.22)  0.68 (0.14)  0.87 (0.25)  1.30 (0.22)  0.56 (0.08)  0.74 (0.08)  0.87 (0.08)  0.61 (0.05)  0.61 (0.04)  0.64 (0.06)  0.58 (0.11)  1.01 (0.15)  1.30 (0.24)  0.56 (0.08)  0.72 (0.12)  1.01 (0.15)  0.49 (0.05)  0.80 (0.12)  0.97 (0.15)  0.43 (0.05)  0.72 (0.06)  0.52 (0.07)  0.68 (0.11)  1.02 (0.10)  0.54 (0.09)  1.18(0.23) 0.78(0.10)  1.17(0.16)  0.87 (0.13) 0.62 (0.06)  0.88 (0.13) 0.59 (0.10)  1.10(0.13) 0.71 (0.11)  0.59 (0.12)  1.00 (0.17)  1.09 (0.16)  0.52 (0.09)  0.49 (0.04)  0.58 (0.07)  0.48 (0.09)  0.79 (0.10)  1.06 (0.15)  0.46 (0.06)  0.49 (0.07)  0.59 (0.08)  0.57 (0.08)  0.83 (0.13)  1.18(0.18)  0.45 (0.05)  0.59 (0.07)  0.60 (0.09)  0.39 (0.06)  0.53 (0.06)  0.75 (0.08)  0.33 (0.06)  0.40 (0.07)  0.37 (0.05)  A C L injured Subjects G M iEMG  125  hypothesize that reduced muscle activation (for the quadriceps) would reduce the extent of tibial motion, the literature is not equivocal. For this reason, we have designed a study that will allow us to explore the interaction. The current research project will examine two research questions: 1) Do changes in knee position affect the amount of quadriceps activation in an ACL deficient limb compared to healthy limbs? 2) Are the changes observed in ACL deficient limbs (ie. reduced quadriceps activation, reduced forces about the knee) exclusive to walking, or can are they observed in other movements? We will measure the muscle activation patterns of various muscles during cycling in two groups: a control group - those individuals with no history of knee injury - and an experimental group - those with a diagnosed unilateral rupture of the ACL. Cycling has been chosen since it is a common activity and one that provides minimal risk to the ACLdeficient individual. Study Procedures After consenting to participate in this study, you will have your leg length and body mass measured. You will then be fitted with bilaterally with surface electrodes to measure muscle activity over the following muscles: rectus femoris, gluteus maximus, vastus lateralis, biceps femoris, semitendinosis, and gastrocnemius. These electrodes will be attached with two-sided adhesive to the skin of your legs and connected to a nearby computer for recording. Prior to applying the electrodes, the skin under the electrodes will be shaved and then lightly abraded to improve electrode contact. An additional ground electrode will be placed on your wrist. Retro-reflective markers will be placed bilaterally over the following bony-landmarks: greater trochanter, lateral knee joint line, lateral malleolus, calcaneus, and lateral aspect of the 5 metatarsal. th  Before the experiment begins, you will be required to complete a stretching warm-up of your own design. During the experimental trials you will be required to cycle for approximately 2 minutes at each of six conditions. The six conditions consist of three power outputs (75 W, 125W, and 175W) at each of two cadences (60 rpm, and 90 rpm). Data will be collected for 15 seconds after steady state has been reached. You will be required to maintain a constant power output and cadence for the duration of each trial. You will be given 3 minutes to rest between conditions. During the trials we will record pedal forces, muscle activity, and we will videotape the movement. The seat height is adjustable and will be fitted to your leg length. This experiment will take about 2 hours to complete.  127  Consent I understand that participation in this study is entirely voluntary and that I may refuse to participate or I may withdraw from the study at any time without any consequences. I also understand that refusal to participate or withdraw from the study at any time will not affect my current or future medical care. I have received a copy of this consent form for my own records. I consent to participate in this study.  Witness Signature  Date  Subject Signature  Date  129  

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