ABSTRACT

ABNORMAL GLUTEAL MUSCLE ACTIVITY AND GROUND REACTION FORCES ASSOCIATED WITH DEVIATIONS IN UNILATERAL TRANSTIBIAL AMPUTEES: A PILOT STUDY

Background: In the United States, a below (BKA) is the most common lower amputation. After a BKA, abnormal gluteal muscle activity during ambulation is directly related to significantly reduced muscle contraction. Individuals with BKA also have abnormal ground reaction forces (GRF) through their prosthetic limb. Objective: The purpose of this study is to examine the differences in gluteal muscle activity and multi-directional GRF during the stance phase of gait in adults with unilateral BKA. Methods: 5 adults with unilateral BKA were recruited for this study. EMG analysis of the bilateral gluteus maximus (GMax) and gluteus medius (GMed) musculature were recorded during MVIC along with GRF during gait trials. Results: Adults with BKA have trends suggesting decreased GMax activity and increased GMed activity during gait. Additionally, there are weightbearing that favor their intact limb. There is a high correlation between muscle activity and asymmetrical weightbearing on the prosthetic limb. Discussion: If there is asymmetrical gluteal muscle strength it can be assumed there will also be asymmetrical GRFs in the anterior-posterior and vertical planes through the prosthetic limb. Key Words: BKA, Ground Reaction Forces, Gluteus Muscle Strength

Melissa Ann Miller May 2017

ABNORMAL GLUTEAL MUSCLE ACTIVITY AND GROUND REACTION FORCES ASSOCIATED WITH GAIT DEVIATIONS IN UNILATERAL TRANSTIBIAL AMPUTEES: A PILOT STUDY

By Melissa Ann Miller

A project submitted in partial fulfillment of the requirements for the degree of Doctor of Physical Therapy in the Department of Physical Therapy College of Health and Human Services California State University, Fresno May 2017 APPROVED For the Department of Physical Therapy:

We, the undersigned, certify that the project of the following student meets the required standards of scholarship, format, and style of the university and the student's graduate degree program for the awarding of the doctoral degree.

Melissa Ann Miller Project Author

Bhupinder Singh (Chair) Physical Therapy

Jennifer Roos Physical Therapy

Jason Schott Prosthetist

For the University Graduate Committee:

Dean, Division of Graduate Studies AUTHORIZATION FOR REPRODUCTION OF DOCTORAL PROJECT

I grant permission for the reproduction of this project in part or in its entirety without further authorization from me, on the condition that the person or agency requesting reproduction absorbs the cost and provides proper acknowledgment of authorship.

X Permission to reproduce this project in part or in its entirety must be obtained from me.

Signature of project author: ACKNOWLEDGMENTS I would like to thank my fiancé, Brad Stillwagon, for being my biggest supporter, inspiration, and motivation to help persons living with amputation. Thank you for showing me that everyone has the chance to achieve their potential in life, and helping me persevere to achieve mine. I would like to thank my family, and my friends in the Class of 2017 for always being so supportive and getting me through the toughest years of my life! Thank you for your love and patience during this time. Finally, I would like to thank my committee members Dr. Bhupinder Singh, PT, PhD Dr. Jennifer Roos, PT, DPT, GCS and Jason Schott, CPO, for their unwavering support, enthusiasm, and assistance with this research project. Thank you for pushing me to pursue my interests, immersing me in the amputee community, and helping me find my calling!

TABLE OF CONTENTS Page

LIST OF TABLES ...... vii

LIST OF FIGURES ...... 1

BACKGROUND ...... 1

Prevalence of Lower Extremity Amputees ...... 1

Muscle Activity During Normal Gait ...... 2

Abnormal Muscle Activity During Gait in Transtibial Amputees ...... 3

Determining Muscle Activity through EMG ...... 4

Ground Reaction Force During Ambulation ...... 4

Gaps in Literature ...... 7

Relevance to Limb Loss Community ...... 8

Purpose ...... 8

METHODS ...... 9

Subjects ...... 9

Procedure ...... 9

Instrumentation ...... 10

Data Analysis ...... 12

RESULTS ...... 14

Subjects ...... 14

EMG Analysis ...... 14

Ground Reaction Force Analysis ...... 15

Analysis of EMG Compared to Ground Reaction Forces ...... 16

DISCUSSION ...... 17

Literature on Strength ...... 19 vi vi Page

Literature on Ground Reaction Force Asymmetry ...... 21

Prosthetic Consideration ...... 23

Clinical Relevance ...... 24

Limitations ...... 25

Additional Research ...... 26

Conclusion ...... 27

REFERENCES ...... 29

TABLES ...... 38

FIGURES ...... 44

APPENDIX: K-LEVELS ...... 50

LIST OF TABLES

Page

Table 1: ICC Table ...... 39

Table 2: Subject Demographics ...... 40

Table 3: MVIC During Manual Muscle Testing ...... 40

Table 4: Average Muscle Activity During Gait Trials ...... 40

Table 5: Percent MVIC During Gait Trials ...... 41

Table 6: Average Forces During Heelstrike ...... 41

Table 7: Average Forces During Midstance ...... 41

Table 8: Average Forces During -Off ...... 42

Table 9: PCC Intact Limb ...... 42

Table 10: PCC Prosthetic Limb ...... 43

LIST OF FIGURES

Page

Figure 1: Prone Gluteus Maximus Testing (knee 90°) ...... 45 Figure 2: Prone Gluteus Maximus Testing (knee 0°) ...... 45 Figure 3: Sidelying Gluteus Medius Testing (knee 0°) ...... 46 Figure 4: Sidelying Gluteus Medius Testing (knee 90°) ...... 46 Figure 5: MVIC ...... 47 Figure 6: Average Muscle Activity During Gait ...... 47 Figure 7: Average GRF Heelstrike ...... 48 Figure 8: Average GRF Midstance ...... 48 Figure 9: Average GRF Toe-Off ...... 49

BACKGROUND

Prevalence of Lower Extremity Amputees There are nearly 2 million people living with limb loss in the United States, with approximately 185 thousand new lower extremity occurring each year.1 Vascular is the most common of amputation, which includes diabetes and peripheral arterial disease, followed by trauma and cancer as second and third most prevalent.1,2 Men are slightly more likely to have a lower limb amputation, with 53.2% incidence in males, and 46.8% incidence in females.3,4 Transtibial amputations, also known as below-knee amputation (BKA) are the most common, representing approximately 65% of lower limb amputations and 71% of dysvascular amputations.2 Current healthcare trends predict a 47% expected increase in BKA from 1995 to 2020.1 Following a BKA, as well as other lower limb amputations, deficits to a person’s neurovascular, neuromuscular, and musculoskeletal systems occur.2 Physical differences in muscle length, strength, (ROM), and limb biomechanics are all changed following disruption of the limbs continuity, and significantly affect a person’s static and dynamic posture.2 The ambulatory abilities of persons with BKA are delineated by their K- level.5 These Medicare adopted code modifiers (K0-K4) classified lower limb amputees based on their functional abilities, as well as their qualification for prosthetics (Appendix). A K-level can be determined through the use of an AMPro Mobility Predictor, which is a valid tool for assessing functional abilities in persons with BKA.6,7 A K3 indicates a person can ambulate at variable cadences within the community, can traverse most environmental barriers, and can have vocational or exercise activity that demands a more advanced prosthetic. A K4 2 2 indicates a person has the ability to perform activities that exceed basic ambulation skills, exhibit high impact stresses, and high energy levels.5

Muscle Activity During Normal Gait The gait cycle of a healthy individual is broken into 2 phases, stance and swing, both initiated by the interactions between the neuromuscular and musculoskeletal systems. Stance is the term used to designate the period of gait in which the is in contact with the ground, beginning with initial contact or strike, and ending with push off. Swing phase is the time in which the entire limb and foot is in the air, beginning with push off, and ending during terminal swing.8,9 Throughout the gait cycle, muscles in synergies to facilitate alternating swing and stance phases for forward propulsion. Agonistic and antagonistic muscles enable variations of lengthening and shortening ratios to allow for the various , knee, and strategies required for gait. At the ankle, muscle activation is phasic, with the plantar flexors activating throughout stance phase, and the dorsiflexors activating during swing phase.8 At the knee, 14 muscles including the quadriceps and hamstrings groups contribute to knee control and activate at integral moments throughout the gait cycle. During stance, the knee extensors act to decelerate knee flexion, and during swing both the knee flexors and extensors contribute to limb progression.9,10 At the hip, gluteus maximus (GMax), gluteus medius (GMed), and gluteus minimus (GMin) facilitate the stance phase of gait. Rectus femoris (RF) and iliopsoas (IP) initiate and control the swing phase.9,10 Additional muscles, including the adductors, are used to control the transitions between swing and stance. The GMax muscle is divided into 2 portions, with the superior half acting as an abductor and the lower half serving as a hip extensor.10,11 At initial contact, 3 3 the superior GMax increases activity to reach 25% effort during single limb loading, facilitating pelvic by the creation of a hip abduction moment. Lower GMax action initiates at the end of terminal swing. Following this activation, the lower GMax decreases its activity level to less than 10% by the end of loading to allow for initiation of hip flexion.11 The GMed and GMin work together during the gait cycle, however analysis of this complex has been limited to the GMed. The Gmed activity begins at the end of terminal swing and increases to a 20% peak through initial contact into mid-stance allowing for contralateral limb swing.9-11

Abnormal Muscle Activity During Gait in Transtibial Amputees of persons with BKA reveals abnormal GMax and GMed muscle activation throughout the gait cycle, more so in stance phase.12 This abnormal gluteal muscle recruitment is greatly affected by the amount of voluntary contraction produced. Decreases in strength significantly reduce static and dynamic hip stability on the prosthetic limb. Isokinetic muscle testing has revealed a significant reduction in strength of the hip stabilizing muscles, more notably the hip extensors and abductors, in lower limb amputees.12 Abnormal recruitment, which is a manifestation of deficits in muscle strength, results in imbalances of the agonist/antagonist ratio, and cause over-reliance on the anterior hip musculature, like the psoas and rectus femoris muscles.12 This asymmetry, along with the reduced hip abductor moment caused by GMed weakness, results in deficits in pelvic symmetry and decreased stride length.13 There is a significant increase in weight distribution on the intact limb compared to the amputated limb during quiet standing (QS), more notably in the anterior and posterior planes.13 Along with pre-existing dynamic postural control 4 4 deficits and core strength deficits, GMed and GMax weakness can also affect pelvic obliquity on the involved limb, further increasing the weightbearing on the intact limb.14,15

Determining Muscle Activity through EMG Surface electromyography (EMG) is a non-invasive method used to quantify the magnitude and timing of muscle activation during various physical tasks, both statically and dynamically.16 These surface electrodes attach directly to the skin over specific points away from tendons or interference from other muscles. The electrical signals produced by the muscles are then transmitted through the surface electrode to a computerized system where clinicians can receive real time accounts of muscle activity, magnitude of contraction, and recruitment patterns both visually and through numerical data.17 Repeatability of EMG data has been established for many isometric activities, including manual muscle testing (MMT).18,19 However, due to its relationship to the skin, interference from other tissue structures can obstruct true results compared to more invasive intramuscular EMG. Despite variance in reliability, it still remains a very accurate and real time analysis of muscular activity in both static and dynamic tasks.16-18

Ground Reaction Force During Ambulation During ambulation, the effects of body weight and force produced during contact with the ground is identified by the ground reaction force (GRF). These forces can exist in 3 planes of motion: anterior-posterior (X), medial-lateral (Y), and vertical (Z). During propulsion, as weight is transmitted through the limb during contact, a force in the ground is created that is of equal magnitude but opposite in direction.9 The anterior-posterior forces are present as the body’s 5 5 center of mass moves over each limb. The medial-lateral forces are affected by the abductor and adductor moments that occur as the body weight is assumed by the stance leg, and the vertical forces increase as gait speed increases.9 A force plate can be used to analyze the differences in GRFs present during foot contact in stance phase. The force plate can then transfer this information to a computerized system, provide visual and numerical expressions of the GRFs through each of the sub-phases within stance.9 Comparison of limbs during static standing or can demonstrate discrepancies in the GRFs or weight acceptance.19 As with EMG, the GRF’s can be affected in those with lower limb amputations. Common practice when collecting GRFs is to normalize them by linearly scaling them to body weight, called division normalization.20 This normalized the peak vertical GRFs with body weight, and allows for better comparison between subjects. This assumes that heavier subjects will produce more GRFs, and those with less body weight will produce less.20 This normalization allows researchers to compare how other contributing factors, such as height, speed, or affect GRFs. Studies have shown little association with body weight and anterior- posterior and medial-lateral GRFs.21 Though this normalization can improve the significance of the results, it does not remove all error.20,21 Muscle activity develops forces and creates a lever that exerts a about 1 or more joints of the body. Dynamic coupling allows for muscle joints to be affected by a single muscle activity. This coupling of muscle synergies allows for the body to accelerate during ambulation, which increased the GRFs.22 A study by Lin et al examined joint forces and how it affected multi-directional GRFs during ambulation, which an inverse dynamic and static optimization relating muscle activity during each phase of the gait cycle.22 This study found that GMax, GMed along with 3 other muscles contributed to vertical support and deceleration 6 6 of the body during stance, affecting the vertical GRFs.22 The quadriceps muscle group directly related to the lateral GRF in the second half of stance phase. GMed opposed the quadriceps muscle and contributed to a medial GRF on the body throughout stance phase.23,24 Additional muscles, like the quadriceps, gastrocnemius, and soleus affected the acceleration of the body, and increased the anterior-posterior GRF in the first half of stance phase, as well as terminal stance.23,24 Throughout the loading response, the magnitude of the vertical GRF exceeds total body weight. In order to accommodate this increased load, the tibialis anterior muscles act as shock absorbers, assisted by eccentric quadriceps activation within 15 to 18 degrees of knee flexion, with the maximum knee- flexion angle reached at foot flat during mid-stance.9 During loading response, the hip extensors which include the GMax, hamstring muscle group, and adductor magnus act as shock absorbers about the hip joint, and prevent further flexion of the hip. Because persons with BKA have compromised anterior tibial and hamstring musculature, as well as a decreased biomechanical advantage of the patellar tendon, there is decreased shock absorption, and disruption of the musculature throughout the hip and knee.25,26 No significance difference was found between GRFs and prosthetic alignment, particularly in those with long established gait parameters.27 Studies have shown differences between weight acceptance and vertical GRFs on the intact limb compared to the prosthetic limb. It has been established that the intact limb has greater weight acceptance during the stance phase of gait, and that the prosthetic limb produces reduced vertical GRFs in the early phases of stance phase.28,29 Studies looking at those with a transfemoral, or above knee amputation 7 7

(AKA) had significantly more weightbearing and more normalized GRFs with a longer residual limb.29

Gaps in Literature Studies on amputee gait have more recently focused on prosthetic biomechanics and kinetics rather than ground reaction forces and muscle activity.30 Contemporary studies have utilized advanced optokinematic technology to look at the relationships of hip and knee moments in persons with BKA, and the relationship of different prosthetic ankle components.31,32 The aim of these studies often focus on observing deviations in joint moments and recommending ways to improve joint mechanics.30,31 Similarly, quantifying the increase in caloric expenditure with persons with unilateral BKA has been a recent topic of research, with studies aiming at examining the increased energy demands.32,33 Multiple studies have been published on the increased caloric expenditure during gait in persons with BKA, and the effects on energy storing prosthetics.32 Though these fields of study are crucial in improving quality of life, decreasing risk of injury and mortality, and improving ambulatory abilities within these individuals, there is a substantial omission of studies examining muscle activity during gait as it related to increased weightbearing or ground reaction forces. Studies examining gait asymmetries and differences focus solely on muscle activity, gait speed, optokinetic observations, or prosthetic type. Muscle activity on the prosthetic limb is assumed to be less than that of the intact limb, however no one study has been effective at examining proximal hip muscle activity in ambulatory persons with BKA as it relates to the 3 planes of GRFs.34-37 8 8 Relevance to Limb Loss Community Understanding the relationship of muscle activity in proximal hip musculature as well as GRF deviations can better predict treatment strategies and ways to improve outcomes in persons with unilateral BKA. Improving an amputee’s proximal hip musculature strength improves gait mechanics, and therefore will limit pain associated with the indicated , improve fit of prosthetic or create a more comfortable prosthetic and then decrease risk of developing osteoarthritis (OA) in their intact limb. Improved fit and decreased pain will significantly improve activity tolerance, willingness to ambulate, and consequently their quality of life.38

Purpose Previous studies have examined gait deviations in amputees, however there is no single study that effectively examines the muscle activity during gait to see how anatomical and biomechanical changes affects gait in persons with BKA. Similarly, no comprehensive study has been conducted on multi-plane GRFs during preferred pace walking in this population. Therefore, the purpose of this study is to examine the differences in gluteal muscle activity and multi-directional GRFs during the stance phase of gait in adults with unilateral BKA. The null hypothesis is that there is no difference between the 2 limbs. The alternative hypothesis is that persons with unilateral BKA will have decreased gluteal muscle activity on the prosthetic limb compared to the intact limb. Additionally, they will have decreased GRFs during stance phase from heel contact to toe off on the prosthetic limb when compared to the intact limb. Finally, there will be a statistically significant correlation between EMG and GRFs on the prosthetic limb.

METHODS

Subjects Prior to collection of data, approval was obtained through the Institutional Review Board of California State University, Fresno. Subjects were sampled from the local amputee community. Inclusion criteria included: age 18-65 years, and possession of a unilateral BKA. Subjects were excluded from the study if they had any of the following: a surgical revision in the last 12 months, any open wounds or sores in the lower extremities, unable to walk without the use of an assistive device, unable to stand for greater than 20 minutes, having poorly controlled cardiovascular disease (PAD, CAD, HTN), have a history of poorly controlled or brittle diabetes, a history of active cancer (or cancer with/without treatment in the last 2 years), and required verbal clearance for physical activity by primary care or orthopedic physician.

Procedure All subjects completed all forms including consent, release of liability, and answered questions based on current medical history. Height, weight, intact limb length, and residual limb length were recorded. Height was measured with a metric ruler against a wall from floor to top of head, with the subject standing flush against the wall barefoot. Weight was measured in Kg on the Kistler© Force Platform with a static stand of 5 seconds on a single force plate. Segmental limb length was measured from center of greater trochanter to the lateral condyle, and distally to the apex of the lateral malleolus or most distal portion of soft tissue. All subject information was entered with a code name, and physical copies of consent forms were kept in a locked cabinet at all times. 10 10

Prior to gait trials on the force platform, the Zenomat was calibrated through Protokinetic© software. Subjects began by walking 3 consecutive passes on the Zenomat pressure mat at their preferred walking pace to extract step and stride length. The Zenomat is similar to the Gaitrite, which is a valid tool for measuring both average and individual step parameters of gait and has excellent reliability.39,40

Instrumentation Wireless EMG electrodes were used to collect muscle activity during maximal volitional isometric contraction (MVIC) and during preferred pace ambulation.41 Skin was prepped for electrode placement by vigorous scrubbing for 15 seconds with an alcohol wipe over the GMax and GMed. Electrode placement was determined based on Rainoldi et al (2004) and followed SENIAM recommendations.42,43 Bipolar surface electrodes (MVAP 1” x 1 7/8”, foam electrode with 2 snaps and 2 gel sites, 1 cm space between electrodes) were used. The sampling rate was set at 1000 Hz utilizing MyoResearch XP™ from Noraxon (Noraxon USA, Scottsdale, Arizona). All raw myoelectric signals were preamplified with an overall gain of 1000. The common rejection ratio rate was set at 100 dB and signal to noise ratio <1 μV RMS baseline noise. The filter to produce a bandwidth was set at 10-1000 Hz. GMax electrode placement was 33.8±11.0 % of the distance from the second sacral vertebra to the greater trochanter, starting from the second sacral vertebra. GMed electrode placement was 33.4±12.8 % of the distance from the iliac crest to the greater trochanter, starting from the greater trochanter.42 Battery packs were connected to each electrode through lead wires, and affixed to the skin 11 11 lateral to the electrodes with double sided tape at a distance to allow for no pull or interference with the electrode pad, clothing, or during requires tasks.43 MVIC was measured based on standardized manual muscle testing positions by Daniels & Worthingham (Figures 1-4).44 MVIC for GMed and GMax was measured in the following positions: sidelying abduction with straight knee, sidelying abduction with bent knee at 30 degrees, and prone hip extension with straight knee, and prone hip extension with knee at 90 degrees. In prone, were placed at 45 degrees flexion with the assistance of 1-2 standard treatment pillows.45 All MVIC testing was conducted against a static gait belt secured to the table to provide an isometric resistance that was consistent throughout all subjects.45 Electrodes placement was maintained throughout the subsequent gait trials. In preparation for collecting data with amputee subjects, an intraclass correlation coefficient (ICC) was carried out with healthy volunteers and revealed an ICC of 0.96, indicating consistency between multiple gait trials on the Kistler© Force Platform (Table 1). Prior to initiation of the gait trials, the force platform was calibrated by InstaCal™ software. The subjects performed 5 trials of walking on the force platform walkway at their preferred pace. Starting position was established based on data collected from the Zenomat, and allowed 3 steps prior to initial contact with the first force plate. Ground reaction force data were obtained using a Kistler© Force Plate (Kistler Instruments, Inc., Amherst, NY). The force plate data were sampled at 300 Hz, and were filtered at 6 Hz. GRF data were run through Bioware, which is a software program that graphs forces produced in the 3 planes of motion: anterior-posterior (X), medial-lateral (Y), and vertical (Z). Numerical representation of these forces in Newtons (N) was taken based on the maximal value at 3 points during ambulation: heel strike, midstance, and toe-off 12 12 on the intact limb as well as the prosthetic limb. These values were ascertained from plotted graphs representing each subject’s average during the subphases of stance phase.

Data Analysis Maximal GRFs during each subphase of stance phase (heel strike, midstance, and toe-off) was calculated for each subject during each trial in the 3 planes of motion (anterior-posterior, medial-lateral, and vertical). These values for the intact limb and the prosthetic limb were compared to one another, as well as averaged across subjects. A 2-tailed T-Test was then performed, which compared the forces produced on the intact limb with those produced on the prosthetic limb. A second 2-tailed T-Test was used to compare the GRFs of the intact limb compared to the prosthetic limb. A 2-tailed T-Test is used if deviations of the estimated limitation are in either direction from a principle value.46 To determine significant difference, an alpha level was set at 0.05. EMG data were exported and analyzed by a single trained researcher. Maximum voluntary isometric contraction (MVIC) for both the intact and prosthetic limb was extracted from the EMG recordings, as well as the mean muscle activity during each trial of ambulation. Mean values were calculated for both the intact and prosthetic limb GMax and GMed in each scenario. In comparing the intact limb and the prosthetic limb during maximal contraction and average muscle activity during ambulation, a 2-tailed T-Test was utilized to determine significance, and the alpha level was set at 0.05.46 A Pearson correlation coefficient (PCC) was calculated comparing the EMG activation during ambulation with the GRFs produced on both the intact and prosthetic limb. A correlation coefficient is a number that quantifies the type of 13 13 correlation and dependence, and weights the statistical relationship between 2 sets of values.47 A numerical value between +1 and -1 is produced, where 1 represents a total positive linear correlation, and 0 is no linear correlation, and -1 is a total negative linear correlation. A high correlation was identified as greater than 0.75.

RESULTS

Subjects Subjects were taken from a sampling of convenience and recruited from local prosthetists, orthotists, and a local amputee support group based on eligibility criteria. A total of 5 subjects were included in the study. The subjects mean ± SD age, height and weight were 51.2±14 years, 178.3±14.3cm, and 94.66±15.29kg, respectively. Average BMI was 30.04±6.1 kg/m2. Average length of residual limb was 73.86±2.52% the intact limb, and the average age of the implant was 16.6±9.44 years. Subject information is presented in Table 2.

EMG Analysis For comparison of both EMG analysis and GRF’s during ambulation between the intact and prosthetic limb, T-Test results are needed. Significance is demonstrated with a value equal to or less than p=0.05. A p-value greater than 0.05 was considered statistically insignificant. Results of the EMG analysis for the intact and prosthetic GMax and GMed can be found in Tables 3-5. Maximal muscle activation during manual muscle testing (MMT) results are represented in Table 3. Average muscle activity for each subjects 3 successful gait trials are represented in Table 4. Average muscle activity compared to MVIC during ambulation for intact GMax was 1.53±1.11%, and 2.34±0.80% for intact GMed. Average percent MVIC during ambulation for the prosthetic GMax was 2.65±1.65%, and 4.28±4.88% for prosthetic GMed, and are represented in Table 5. Visual representation of muscle activity can be seen in Figures 5-6. T-Test results of the EMG recordings of the GMax and GMed during MMT demonstrate no statistical significance between the intact limb compared to the prosthetic limb. T-Test results of the intact and prosthetic GMax and GMed were 15 15

0.18 and 0.20, respectively. During ambulation, a T-Test of average muscle activity demonstrated no statistical significance between the intact and prosthetic limbs, with a GMax and GMed having a value of 0.41 and 0.21, respectively. It should be noted that the GMed on the prosthetic limb had a relatively higher activation rate compared to the intact GMed, although did not reach clinically significant peak forces.

Ground Reaction Force Analysis Three complete gait trials were evaluated for each subject, which excluded trials in which there was an absence of force plate contact or trials in which full foot contact was not achieved. Averages of the 3 complete gait trials were obtained for each subject in all planes of motion within the 3 main points of contact during gait. Averages were calculated for all subjects for both the intact limb and the prosthetic limb. Average forces during heel strike, midstance, and toe-off for each subject’s intact and prosthetic limb are represented in Tables 6-8. Visual representation of peak GRFs for each plane of motion during each sub-phase of stance for the sampling of subjects is present in Figures 7-9. During heel strike, the T-Test revealed significance in the anterior-posterior plane (0.001), and no statistical significance in the medial-lateral plane (0.523) and vertical plane (0.179). During midstance, there was no statistical significance in any of the 3 planes, with the p-values of the anterior-posterior, medial-lateral, and vertical planes equaling 0.775, 0.9, and 0.369, respectively. During toe-off, statistical significance was found in the anterior posterior plane (0.015), and no significance in the medial lateral and vertical planes, 0.484 and 0.365, respectively. While no clinical significance can be confidently confirmed from 16 16 these results, visual representation of peak GRFs demonstrated in the figures indicate the relative consistency that intact limb GRFs are higher than that of the prosthetic limb during all sub-phases of stance phase in all planes of GRFs.

Analysis of EMG Compared to Ground Reaction Forces While EMG data did not show any statistical significance, and only moderate significance was found with GRFs, a PCC was used to examine the correlation between the gluteal muscle activity on the intact and prosthetic limb, and the GRFs associated with each subphase of stance. The results of this correlation are presented in Tables 9-10. At heel strike, there was a high statistically significant correlation between the prosthetic limb’s GMax (0.949) and GMed (0.931) activity and the GRFs in the medial-lateral plane on the prosthetic limb. Additionally, there was also a significant correlation between the prosthetic limb’s GMax (0.832) and GMed (0.904) activity and the GRFs in the vertical plane on the intact limb. At midstance, there was a significant correlation between the prosthetic limb’s GMax (0.777, 0.851) and GMed (0.783,0.863) muscle activity and the anterior-posterior and medial-lateral GRFs, respectfully. At toe-off, a high statistical correlation was found between the intact limb’s GMax (0.79) and prosthetic limb’s GMax (0.79), and the vertical GRFs through the intact limb. There was no other statistically significant correlation found between muscle activity on the intact limb, and GRFs in any of the planes of motion.

DISCUSSION

The purpose of this study was to determine if there were significant muscle activation and ground reaction force asymmetries in adults with unilateral BKA during preferred walking pace ambulation. It can be concluded that the null hypothesis was partially accepted, as there was little statistically significant differences found in EMG and GRFs results between limbs. The alternative hypothesis was also partially accepted, as there were some clinically significant differences in GRFs during heel strike, and trends of additional differences during toe off. Additionally there was a high statistically significant correlation between GMax and GMed activity and medial-lateral and vertical GRFs. Additional trend analysis suggests differences in muscle activity, though more in-depth analysis must be performed to determine the extent of these results. Results of the EMG study demonstrated no statistical significance, however did demonstrate potential trends. Prosthetic GMax demonstrated the lowest MVIC and least activity during ambulation. However, the prosthetic GMed demonstrated significantly higher MVIC and activation during ambulation. Increased average prosthetic GMed activity could be attributed to the increased activity even during rest. During ambulation, the prosthetic GMed maintained a higher activity level than the intact GMed. Even during swing phase, the prosthetic GMed maintained higher activity. Because EMG recordings were not separated by phase of gait (stance or swing), true gluteal muscle activity in stance phase could not be gathered. The over-activity of the prosthetic limb’s GMed could be due to the abnormal GMax activity. In normal gait, the muscles work in synergy to provide vertical support.22 It could also be hypothesized that there were interferences from other muscle groups, or from the clothing over the electrode. Because activity of 18 18 the musculature in close proximity of the GMed was not collected, true GMed activity could not be inferred. While there was only significance in the anterior-posterior plane during heel strike (0.001) and toe-off (0.015), plotting the results demonstrated a general trend that the intact limb produces more ground reaction forces in during every phase of gait. One plausible explanation for the differences in GRF’s is that the biomechanics of the prosthetic limb may absorb forces at heel strike, and release it at toe-off.48 During toe-off, the prosthetic foot’s rigid design could account for the decreased vertical GRFs. Because the prosthetic foot is more rigid, and it’s intent to store energy at heel strike, it dampens the vertical forces until toe-off.48,49 While prosthetics try to mimic natural foot motion, the intact foot’s pliability allows for improved distribution of GRFs more effectively than the prosthetic limb.49 During midstance, the body’s proportional progression over each limb must demonstrate relative equality, otherwise no forward progression would be created.50 While comparing the insignificance of the EMG data and the variable significance of the GRFs, a PCC of the values determined that there was a significant correlation with Gmax and Gmed muscle activity and ground reaction forces in the medial-lateral plane at heel strike, and in the anterior-posterior and medial-lateral planes in midstance. On the prosthetic limb at heel strike, the GMax and GMed force production correlated with medial-lateral GRFs, indicating that muscle activity affects the weightbearing abilities and forces produced during limb contact. Similarly, this muscle activity also correlated with anterior-posterior and medial-lateral force production in midstance. The high correlation of medial- lateral forces in midstance with GMed activity can be directly related to the abductor moments about the hip, and the relationship the hip abductors and adductors play in controlling deviations within this plane of motion.11 At heel 19 19 strike, the GMed activity on the prosthetic limb also has a positive correlation with the vertical ground reaction forces on the intact limb. Results of the PCC could indicate a relationship between the decreased GMax and increased GMed activity and the weightbearing asymmetries present in persons with unilateral BKA. It could be hypothesized that the prosthetic limb GMed has to activate more throughout the stance phase of gait and during the initiation of swing phase to elevate the limb and maintain pelvic symmetry. While the results do not imply causation, and should not be interpreted as such, these results can be used to infer that if a person with unilateral BKA presents with abnormal GMax or GMed activity, then there is high statistical probability that abnormal GRFs will be present, and conversely abnormal GRFs will present with abnormal muscle activity. The low correlation between the intact GMax and GMed and forces produced on the intact and prosthetic limb demonstrate that any abnormalities present in either of these variables have any relationship to one another.

Literature on Strength Asymmetry While this study did not identify any significant muscle activity differences between the intact and prosthetic limb, other studies have found significant differences. One study by Isakov et al found a significant (P<0.05) increased biceps femoris activity during the stance phase of gait in order to improve stability on the prosthetic limb.51 Studies comparing highly active and inactive transtibial amputees proximal hip strength with concentric and eccentric strength testing demonstrated differences based on the activity level of the population. While there were little strength asymmetries in the highly active group, there were significant asymmetries in the inactive group.51 A study by Nolan et al also found that highly 20 20 active persons with BKA produced greater MVIC of their prosthetic limb.52 While activity level was not recorded in our study, this could explain the lack of statistical significance between limbs recording muscle activity. Based on observations, and ability to perform functional tasks post gait trials (sit to stand, single limb , static standing) all subjects appeared to be K3-K4, and highly active in either vocational or recreational activities. In the study by Nolan, “highly active” was defined as those who exercised at least 3 times per week for recreational purposes. Sedentary or “inactive” subjects were defined as individuals who were only physically active when performing ADLs.52 This inconsistency within our study could be related to our lack of control of activity level. Inclusion of a fitness questionnaire that assessed fitness and activity level may be appropriate to better determine how physical fitness affects muscle activity. Additional studies consistently find high statistical differences (P<0.001) in the strength asymmetry between the quadriceps and hamstring muscle group of the prosthetic limb compared to the intact limb.53 A similar study looking at 8 unilateral transtibial and 8 control subjects assessed gait variables related to OA risk factors. These included knee external adduction moment, knee adduction moment load rate, vertical ground reaction force load rate, and 3 strength measures. Four of the 6 variables were more asymmetrical in the amputee group compared to the control group (p<0.05), with knee extension strength asymmetry being directly related to knee adduction moment load rate asymmetry, and knee flexion strength asymmetry was moderately related to the vertical ground reaction force on the intact limb. This strength asymmetry has a moderate correlation to increased deterioration of the joint, and increased OA risk factors.38 While knee adduction moments were not 21 21 calculated as part of our study, trends of higher vertical GRFs in the anterior- posterior direction indicate higher stresses on the knee joint of the amputee limb.

Literature on Ground Reaction Force Asymmetry It has been expressed in the literature that persons with BKA stand with more weight on the intact limb compared to the prosthetic limb.26,28 Nadollek et al examined QS in adults with unilateral BKA and found increased anterior-posterior sway in the intact limb compared to the prosthetic limb.13 Our study found that during ambulation, mid-stance demonstrated the least amount of variance in the medial-lateral plane, which is expected due to the normalized full foot contact. Heel strike and toe-off both had marked differences, with toe-off demonstrating the greatest difference in the medial-lateral plane, with the intact limb having greater excursion. In our study, while increased sway was expected, we found little difference from the prosthetic limb compared to the intact limb. The lack of significance found in our study may be due to all of our subjects demonstrating the appropriate 1/3 length of their intact. A longer residual limb can normalize GRFs and weightbearing through the prosthetic limb.28 While no notable differences were noted in the medial-lateral plane throughout the stance phases, we did see variance in the anterior-posterior plane at both heel strike and toe-off. As subjects center of mass displaced forward during gait, the intact limb demonstrated greater anterior-posterior GRFs due to normal foot mechanics.54 In normal gait, as the foot accepts weight, it transitions from a supinated position to midfoot pronation and finally ends with greatest force production over the great toe. The intact foots pliability and ability to react to ground reaction forces allows this anterior-posterior movement.54 On the prosthetic limb, while multi-axial feet assist in this transition, the fluidity can be 22 22 compromised.55-56 While this study looked at preferred pace walking, a study by Baum et al examined ground reaction forces in submaximal . Researchers found that as amputees spend less time on their prosthetic limb, and more time on their intact limb, there is a quicker transfer onto the intact limb, promoting an anterior and lateral force vector about the intact foot expressed as more movement.26 Numerous studies have been done on vertical GRF, and while our study did not show clinical significance, the trends associated with peak GRFs support the evidence that amputees have decreased vertical GRFs on the prosthetic limb compared to the intact limb (Tables 3-5).57,58 Our study also examines the relationship of all force vectors, and how their combined effects impact amputee gait. Based on trends noted in peak forces, as the prosthetic limb lands during ambulation, it produced less vertical force, less anterior-posterior force, and similar medial-lateral forces as compared to the intact limb. With increased vertical loading on the intact limb, comes increased force distributed through the joints. Previous studies done on numerous gait deviations explain that these increased forces lead to accelerated deterioration of the intact hip and knee joints.58 Similarly, biomechanical compensations for the prosthetic limb and its decreased weight acceptance are present up the chain with varying postural deficits, trunk leans, pelvic motion, and even excessive arm swings.14,15 Additional factors that contribute to increased loading through a limb include stride length, weight, and BMI. Noted by our initial data found through Zenomat testing, amputees have a decreased stride length on the prosthetic limb compared to the intact limb. A greater stride length leads to more force production at heel strike, as it takes greater muscular force to propel the body forward. Smaller steps lead to decreased force production.9 23 23

Research demonstrates increased vertical GRFs as weight and BMI increase. While BMI was not controlled for, it may account for the differences in vertical ground reaction forces noted in those with a greater BMI. A study by Pamyukoff et al found that individuals classified as obese (BMI >30) demonstrated a higher heel strike transient (HST) than those with a normal BMI (25 kg/m2).59 The HST is a rapid, transient rise in the vertical ground reaction forces following heel strike, and is indicative of poorly controlled loading. The researchers noted that obese individuals experience greater loading rates due to poor eccentric quadriceps control and quadriceps weakness.59 As persons with BKA can demonstrate deficiencies in eccentric quadriceps control, a similar HST may occur. Pamyukoff et al also found that there was no difference between the standardized speed (1.0 m/s) or walking at a self-selected pace in the obese group.60 Additionally study by Browning and Kram indicated that obese adults walk 0.3 m/s slower than their normal weight counterparts.61 These 2 studies support the decision to have the subjects included in this study walk at a self- selected pace, as there would be no difference if it was standardized under 1.0 m/s.

Prosthetic Consideration More current research studies focus on the correlation of prosthetic type and muscle activity during gait. Prosthetics that include energy storage and return (ESAR) feet have been developed to provide enhanced function by storing and returning mechanical energy through elastic structures. In one study a total of 5 foot conditions were analyzed: solid ankle (SA), stiff forward-facing ankle (FA), compliant FA, stiff reverse-facing ankle (RA) and compliant RA. The ESAR decreased the activity of muscles that contribute to body forward propulsion and increased the activity of muscles that provide body support.62 The 24 24 compliant ankles generally caused a greater change in muscle activity than the stiff ankles, but without a corresponding increase in energy return. Ankle orientation also had an effect, with RA generally causing a lower change in muscle activity than FA. These results highlight the influence of ESAR stiffness on muscle activity and the importance of prescribing appropriate prosthetic foot stiffness to improve ambulatory outcomes.62

Clinical Relevance The findings of this study add to the literature regarding gait asymmetries in adults with unilateral BKA. Studies examined the relationship of proximal hip and knee strength, and found that these strength deficits had a moderate relationship between increased weightbearing on the intact limb, and therefore increased risk of developing OA on the intact limb.38,63 When examining gait in a person with BKA, it should be understood that the gait deviations noted in this study can result in anatomical and biomechanical changes that decrease efficiency of amputee gait and cause decreased efficiency during ambulation and increased caloric expenditure. Understanding the discrepancies between the intact and prosthetic limb, and hypothesizing what secondary problems could occur and how to remedy them can aid clinicians in improving these asymmetries to decrease energy demands and improve efficiency during ambulation.64,65 The high statistical correlation found between GMax and GMed muscle activity and ground reaction forces can be used as a tool by clinicians to determine impairments present in BKAs. Because of the high statistical significance, it can be proposed that if clinicians find GMax or GMed strength or activity deficits on the prosthetic limb, they can assume that there is decreased GRF’s on the prosthetic limb. This may allow clinicians to work on weightbearing asymmetries 25 25 through normalizing proximal hip strength without the need to confirm these discrepancies with a force platform or expensive equipment.

Limitations While this study added to the literature supporting differences in muscle activity and GRF’s between the intact limb and prosthetic limb in adults with unilateral BKA, there were substantial limitations throughout the study. There was a short amount of time allotted for data collection, and a small sample size of participants. The PCC was only correlated with muscle activity Additionally, there was a disproportionate lack of female representation in the study, with current statistics from the local Fresno Amputee Coalition depicting a 1:2 female to male ratio.66 Gait speed was not controlled for, however as previously discussed, may not have contributed to increased vertical GRFs in preferred pace walking in obese individuals.59 A study by Keller et al did find vertical ground reaction forces did increase linearly as gait speed increased, however remained constant at higher speeds.67 One notable limitation is the relationship that the we did not examine the subjects limb dominance, and how potential amputation of their dominant leg affected it’s force production. Wang and Watanabe studied vertical GRFs and found dominant limbs produce more force.68 Additional limitations included the lack of control for variations in prosthetic type as well as the age of prosthetic. Energy storing prosthetics affect vertical GRFs, and may have been normalized based on industry data.37 Time since last visit with prosthetists/orthotist and prosthetic adjustment was also not controlled for, as some subjects had not had a recent revision and an older prosthetic that was not functioning optimally and exacerbating gait deviations. Weight was also not controlled for, which greatly 26 26 affected the force produced during ambulation and did not include division normalization, though may have varied in reliability. A BMI above 30 can have an effect on vertical GRFs up to point in which they plateau.59,69 True muscle activity and relationships between lower limb synergies was compromised as only 4 electrodes were used, and no distal or lower limb muscle activity was recorded. Additionally, results of the PCC indicate correlation between muscle activity during both the stance and swing phase of gait, as muscle activity during these individual phases was not controlled for. While there was significant correlation between the 2 variables and it could be hypothesized that controlling for activity during stance phase alone may wield better statistical significance, a true correlation cannot be assumed.

Additional Research Additional research is needed to improve the significance of this study. A larger sample size would improve the impact and relevance of the results. Additional female subjects and those under the age of 45 would also be beneficial in controlling for gender and any differences noted with age.70 Controlling for activity level and grouping subjects based on their current activity level could positively affect outcomes on muscle activity. As noted in the study by Nolan on muscle activity and activity level, examining lower activity level subjects could better identify asymmetries between the intact and prosthetic limb.52 Additional research examining the differences in functional tasks, include static standing, transfers, and squatting is recommended as these can be excellent tools for examining disproportionate weight distribution throughout the intact and prosthetic limb. Similar to studies conducted by Nadollek et al and Lenka and Tiberwala should be performed to add to the literature relating structural changes 27 27 and QS abilities in adults with unilateral BKA.13, 70 Further investigation of QS abilities as it relates to ambulatory abilities would be beneficial in truly identifying the repercussions of weightbearing asymmetries in this population. Additionally, as abnormal anterior and posterior proximal hip and knee muscular activity has been observed in adults with unilateral BKA, future studies should include EMG recordings of the anterior hip, quadriceps, and hamstring muscle groups. Once appropriate results have been obtained for all muscle synergies, and a comprehensive knowledge of the relationship of lower extremity musculature and its effects on weightbearing is obtained, additional research can address the muscle activity deficits and determine effective and appropriately dosed exercise to best improve these asymmetries.

Conclusion Results of this study conclude that adults with unilateral BKA have weightbearing asymmetries that favor their intact limb. There were statistically significant differences during heel strike and toe-off in the anterior-posterior GRFs. While it is understood that weightbearing asymmetries can exist in many populations and even healthy individuals may never be truly uniform, clinicians can use this information when examining amputee gait to assist in creating amputee rehabilitation programs designed to improve muscle strength imbalances and weightbearing symmetry. The results from the PCC enable clinicians to make assumptions based on the high correlation between the GMax and GMed activity and the GRFs on the prosthetic limb, and predict the presence of these differences. These findings may inform prospective researchers on the existing asymmetries in muscle activity and GRFs present in persons with unilateral BKA. Further 28 28 research should be conducted to examine these differences and add to the statistical significance of these findings.

REFERENCES

REFERENCES

1. Ziegler‐ GrahamK, MacKenzie E, Ephraim P, Travison T, Brookmeyer R. Estimating the Prevalence of Limb Loss in the United States: 2005-2050. Archives of Physical Medicine and Rehabilitation. 2008;89(3)(422-9).

2. Most RS, Sinnock P. The Epidemiology of Lower Extremity Amputations in Diabetic Individuals. Diabetes Care. 1983;6(1):87-91. doi:10.2337/diacare.6.1.87.

3. Akiode O, Shonubi AMO, Musa A, Sule G. Major limb amputations: an audit of indications in a suburban surgical practice. Journal of the National Medical Association. 2005;97(1):74-78.

4. Robbins JM, Strauss G, Aron D, Long J, Kuba J, Kaplan Y. Mortality Rates and Diabetic Foot Ulcers. Journal of the American Podiatric Medical Association. 2008;98(6):489-493. doi:10.7547/0980489.

5. Gailey R, Roach K, Applegate E. The Amputee Mobility Predictor: An instrument to assess determinants of the lower-limb amputee's ability to ambulate. Archives of Physical Medicine and Rehabilitation. 2002;83(5):613-627. doi:10.1053/apmr.2002.32309.

6. Gailey R. Predictive Outcome Measures Versus Functional Outcome Measures in the Lower Limb Amputee. JPO Journal of Prosthetics and Orthotics. 2006;18(Proceedings):P51-P60. doi:10.1097/00008526- 200601001-00006.

7. Resnik LBorgia M. Reliability of Outcome Measures for People With Lower-Limb Amputations: Distinguishing True Change From Statistical Error. Physical Therapy. 2011;91(4):555-565. doi:10.2522/ptj.20100287.

8. Boakes JL, Rab GT. “Muscle activity during walking.” Human Walking. Lippincott Williams and Wilkins, Baltimore (2006).Available from: SPORTDiscus with Full Text. Accessed November 06, 2016.

9. Perry J, Burnfield, J. “Gait Analysis: normal and pathological function.”.(1992): 271-279.

10. Nielsen JB. How we Walk: Central Control of Muscle Activity during Human Walking. The Neuroscientist. 2003;9(3):195-204. doi:10.1177/1073858403009003012. 31 31 11. Ivanenko YP, Poppele RE, Lacquaniti F. Five basic muscle activation patterns account for muscle activity during human locomotion. The Journal of Physiology. 2004;556(1):267-282. doi:10.1113/jphysiol.2003.057174.

12. Croisier J, Maertens De Noordhout B, Maquet D. Isokinetic evaluation of hip strength muscle groups in unilateral lower limb amputees. Orbiulgacbe. 2017. Available at: http://orbi.ulg.ac.be/handle/2268/12006. Accessed November 19, 2016.

13. Nadollek H, Brauer S, Isles R. Outcomes after trans-tibial amputation: the relationship between quiet stance ability, strength of hip abductor muscles and gait. Physiotherapy Research International. 2002;7(4):203-214. doi:10.1002/pri.260.

14. Molina-Rueda F, Alguacil-Diego IM, Cuesta-Gómez A, Iglesias-Giménez J, Martín-Vivaldi A, Miangolarra-Page JC. , and hip pattern in the frontal plane during walking in unilateral transtibial amputees: biomechanical analysis. Brazilian Journal of Physical Therapy. 2014;18(3):252-258. doi:10.1590/bjpt-rbf.2014.0032.

15. Rueda FM, Diego IMA, Sánchez AM, Tejada MC, Montero FMR, Page JCM. Knee and hip internal moments and upper-body kinematics in the frontal plane in unilateral transtibial amputees. Gait & Posture. 2013;37(3):436-439. doi:10.1016/j.gaitpost.2012.08.019.

16. Christ CB, Slaughter MH, Stillman RJ, Cameron J, Boileau RA. Reliability of Select Parameters of Isometric Muscle Function Associated With Testing 3 Days × 3 Trials in Women. Journal of Strength and Conditioning Research. 1994;8(2):65-71. doi:10.1519/00124278-199405000-00001.

17. Kellis E, Katis A. Reliability of EMG -spectrum and amplitude of the semitendinosus and biceps femoris muscles during ramp isometric contractions. Journal of Electromyography and Kinesiology. 2008;18(3):351-358. doi:10.1016/j.jelekin.2006.12.001.

18. Sleivert, GG and Wenger, HA. Reliability of measuring isometric and isokinetic peak torque, rate of torque development, integrated electromyography, and tibial nerve conduction velocity. Arch Phys Med Rehab 75: 1315-1321, 1994. Available from: CINAHL Plus with Full Text, Ipswich, MA. Accessed December 22, 2016. 32 32 19. Inman V, Eberhart H. The Major Determinants in Normal and Pathological Gait. The Journal of Bone & Joint Surgery. 1953;35(3):543-558. doi:10.2106/00004623-195335030-00003.

20. Wannop J, Worobets J, Stefanyshyn D. Normalization of Ground Reaction Forces, Joint Moments, and Free Moments in Human Locomotion. Journal of Applied Biomechanics. 2012;28(6):665-676. doi:10.1123/jab.28.6.665.

21. Bates B, Osternig L, Sawhill J, James S. An assessment of subject variability, subject-shoe interaction, and the evaluation of running shoes using ground reaction force data. Journal of Biomechanics. 1983;16(3):181-191. doi:10.1016/0021-9290(83)90125-2.

22. Lin YC, Kim, HJ, Pandy MG. A computationally efficient method for assessing muscle function during human locomotion. International Journal for Numerical Methods in . 2011; 27: 436–449. doi:10.1002/cnm.1396.

23. Anderson FC, Pandy M. Individual muscle contributions to support in normal walking. Gait & Posture. 2003;17(2):159-169. doi:10.1016/s0966- 6362(02)00073-5.

24. Carley K, Dart K, Vos M. Normative Database of Adult Unilateral Trans- Tibial Amputee Gait. Grand Valley State University 1997. Available at: http://fliphtml5.com/viso/kfuc/basic/51-98. Accessed December 22, 2016.

25. Raya M, Gailey R, Fiebert I, Roach K. Impairment variables predicting activity limitation in individuals with lower limb amputation. Prosthetics and Orthotics International. 2010;Vol.34(1), P.73-84, 34(1), 73-84. doi: 10.3109/03093640903585008

26. Baum B, Hobara H, Kim Y, Shim J. Amputee Locomotion: Ground Reaction Forces during Submaximal Running with Running-Specific Prostheses. Journal of Applied Biomechanics. 2016;32(3):287-294. doi:10.1123/jab.2014-0290.

27. Jonkergouw N, Prins MR, Buis AWP, Wurff PVD. The Effect of Alignment Changes on Unilateral Transtibial Amputee’s Gait: A Systematic Review. Plos One. 2016;11(12). doi:10.1371/journal.pone.0167466.

28. Baum B, Schnall B, Tis J, Lipton J. Correlation of residual limb length and gait parameters in amputees. Injury. 2008;39(7):728-733. doi:10.1016/j.injury.2007.11.021. 33 33 29. Eshraghi A, Osman NAA, Karimi M, Gholizadeh H, Soodmand E, Abas WABW. Gait Biomechanics of Individuals with Transtibial Amputation: Effect of Suspension System. PLoS ONE. 2014;9(5). doi:10.1371/journal.pone.0096988.

30. Powers CM, Rao S, Perry J. Knee kinetics in trans-tibial amputee gait. Gait & Posture. 1998;8(1):1-7. doi:10.1016/s0966-6362(98)00016-2.

31. Herdiman L, Adiputra IN, Tirtayasa K, Manuaba IBA. Improvement in walking efficiency of transtibial amputee using prosthetic leg with multi- axis joint and energy store return ankle. Proceedings of the Joint International Conference on Electric Vehicular Technology and Industrial, Mechanical, Electrical and Chemical Engineering (ICEVT & IMECE). 2015. doi:10.1109/icevtimece.2015.7496644.

32. Hafner B, Sanders J, Czerniecki J, Fergason J. Energy storage and return prostheses: does patient perception correlate with biomechanical analysis?. Clinical Biomechanics. 2002;17(5):325-344. doi:10.1016/s0268- 0033(02)00020-7.

33. Gailey RS, Wenger MA, Raya M, Kirk N, Erbs K, Spyropoulos P, Nash MS. Energy expenditure of trans-tibial amputees during ambulation at self- selected pace. Journal of Prosthet and Orthotics. 1994; 18(2): 84–91. doi:10.3109/03093649409164389.

34. Gard S. Use of Quantitative Gait Analysis for the Evaluation of Prosthetic Walking Performance. JPO Journal of Prosthetics and Orthotics. 2006;18(Proceedings):P93-P104. doi:10.1097/00008526-200601001- 00011.

35. Nolan L, Wit A, Dudziñski K, Lees A, Lake M, Wychowañski M. Adjustments in gait symmetry with walking speed in trans-femoral and trans-tibial amputees. Gait & Posture. 2003;17(2):142-151. doi:10.1016/s0966-6362(02)00066-8.

36. Isakov E, Keren O, Benjuya N. Trantibial amputee gait: Timedistance parameters and EMG activity. Prosthetics and Orthotics International. 2000;24(3):216-220. doi:10.1080/03093640008726550. 34 34 37. Zmitrewicz R, Neptune R, Walden J, Rogers W, Bosker G. The Effect of Foot and Ankle Prosthetic Components on Braking and Propulsive Impulses During Transtibial Amputee Gait. Archives of Physical Medicine and Rehabilitation. 2006;87(10):1334-1339. doi:10.1016/j.apmr.2006.06.013.

38. Lloyd C, Stanhope S, Davis I, Royer T. Strength asymmetry and osteoarthritis risk factors in unilateral trans-tibial, amputee gait. Gait & Posture. 2010;32(3):296-300. doi:10.1016/j.gaitpost.2010.05.003.

39. Webster K, Wittwer J, Feller J. Validity of the GAITRite® walkway system for the measurement of averaged and individual step parameters of gait. Gait & Posture. 2005;22(4):317-321. doi:10.1016/j.gaitpost.2004.10.005.

40. Menz H, Latt M, Tiedemann A, Mun San Kwan M, Lord S. Reliability of the GAITRite® walkway system for the quantification of temporo-spatial parameters of gait in young and older people. Gait & Posture. 2004;20(1):20-25. doi:10.1016/s0966-6362(03)00068-7.

41. Kollmitzer J, Gerold ER, Kopf A. Reliability of surface electromyographic measurements. Clinical Neurophysiology. 1999;110(4):725-734. doi:10.1016/s1388-2457(98)00050-9.

42. Rainoldi A, Melchiorri G, Caruso I. A method for positioning electrodes during surface EMG recordings in lower limb muscles. Journal of Neuroscience Methods. 2004;134(1):37-43. doi:10.1016/j.jneumeth.2003.10.014.

43. Electrode Placement Guidelines: Gluteus Medius and Gluteus Maximus. Seniam.org. Accessed: November 01, 2016. http://www.seniam.org/.

44. Hislop H, Avers D, Brown M.. Daniel's and Worthingham's Muscle Testing Techniques Of Manual Examination and Performance Testing. 5th ed. Philadelphia: WB Saunders; 1986:182-183.

45. Robertson J. F. P. Kendall and E. K. McCreary "Muscles, Testing and Function" (Third Edition). British Journal of Sports Medicine. 1984;18(1):25-25. doi:10.1136/bjsm.18.1.25.

46. Meng X, Rosenthal R, Rubin D. Comparing correlated correlation coefficients. Psychological Bulletin. 1992;111(1):172-175. doi:10.1037//0033-2909.111.1.172. 35 35 47. Katz., Mitchell H. (2006) Multivariable Analysis – A Practical Guide for Clinicians. 2nd Edition. Cambridge University Press. ISBN 978-0-521- 54985-1. ISBN 0-521-54985-X doi:10.2277/052154985X

48. Klodd E, Hansen A, Fatone S, Edwards M. Effects of prosthetic foot forefoot flexibility on gait of unilateral transtibial users. The Journal of Rehabilitation Research and Development. 2010;47(9):899. doi:10.1682/jrrd.2009.10.0166.

49. Powers CM, Torburn L, Perry J, Ayyappa E. Influence of prosthetic foot design on sound limb loading in adults with unilateral below-knee amputations. Arch Physical Medicine Rehabilitation. 1994;75(7):825-29. [PMID: 8024435]

50. Isakov, H. Burger, J. Krajnik, M. G E. knee muscle activity during ambulation of trans-tibial amputees. Journal of Rehabilitation Medicine. 2001;33(5):196-199. doi:10.1080/165019701750419572.

51. Nolan L. Lower Limb Strength in Sports-Active Transtibial Amputees. Prosthetics and Orthotics International. 2009;33(3):230-241. doi:10.1080/03093640903082118.

52. Tugcu I, Yilmaz B, Goktepe A, Taskaynatan M, Yazicioglu K. Muscle Strength and Bone Mineral Density in Mine Victims with Transtibial Amputations. International Journal of Prosthetics and Orthotics. 2009;33(4):299-306. doi: 10.3109/03093640903214075.

53. Dixon Pearsall D. Gait Dynamics on a Cross-Slope Walking Surface. Journal of Applied Biomechanics. 2010;26(1):17-25. doi:10.1123/jab.26.1.17.

54. Enoka R, Miller D, Burgess E. Below-knee Amputee Running Gait. American Journal of Physical Medicine & Rehabilitation. 1982;61(2):66???84. doi:10.1097/00002060-198204000-00002.

55. Tominaga S, Sakuraba K, Usui F. The effects of changes in the sagittal plane alignment of running-specific transtibial prostheses on ground reaction forces. Journal of Physical Therapy Science. 2015;27(5):1347- 1351. doi:10.1589/jpts.27.1347.

56. Bolger D, Ting L, Sawers A. Individuals with transtibial limb loss use interlimb force asymmetries to maintain multi-directional reactive balance control. Clinical Biomechanics. 2014;29(9):1039-1047. doi:10.1016/j.clinbiomech.2014.08.007. 36 36 57. Gailey R. Review of secondary physical conditions associated with lower- limb amputation and long-term prosthesis use. The Journal of Rehabilitation Research and Development. 2008;45(1):15-30. doi:10.1682/jrrd.2006.11.0147.

58. Pamukoff D, Dudley R, Vakula M, Blackburn J. An evaluation of the heel strike transient in obese young adults during walking gait. Gait & Posture. 2016;49:181-183. doi:10.1016/j.gaitpost.2016.07.001.

59. Pamukoff D, Lewek M, Blackburn J. Greater vertical loading rate in obese compared to normal weight young adults. Clinical Biomechanics. 2016;33:61-65. doi:10.1016/j.clinbiomech.2016.02.007.

60. Browning R, Kram R. Effects of Obesity on the Biomechanics of Walking at Different Speeds. Medicine & Science in Sports & Exercise. 2007;39(9):1632-1641. doi:10.1249/mss.0b013e318076b54b.

61. Ventura J, Klute G, Neptune R. The effect of prosthetic ankle energy storage and return properties on muscle activity in below-knee amputee walking. Gait & Posture. 2011;33(2):220-226. doi:10.1016/j.gaitpost.2010.11.009.

62. Struyf P, van Heugten C, Hitters M, Smeets R. The Prevalence of Osteoarthritis of the Intact Hip and Knee Among Traumatic Leg Amputees. Archives of Physical Medicine and Rehabilitation. 2009;90(3):440-446. doi:10.1016/j.apmr.2008.08.220.

63. Waters R, Perry J, Antonelli D, Hislop H. Energy cost of walking of amputees. The Journal of Bone & Joint Surgery. 1976;58(1):42-46. doi:10.2106/00004623-197658010-00007.

64. Zemke L. Fresno Amputee Coalition: Population of Central Valley Lower Limb Amputees. 2017.

65. Waters RMulroy S. The energy expenditure of normal and pathologic gait. Gait & Posture. 1999;9(3):207-231. doi:10.1016/s0966-6362(99)00009-0.

66. Keller T, Weisberger A, Ray J, Hasan S, Shiavi R, Spengler D. Relationship between vertical ground reaction force and speed during walking, slow jogging, and running. Clinical Biomechanics. 1996;11(5):253-259. doi:10.1016/0268-0033(95)00068-2. 37 37 67. Wang Y, Watanabe K. Limb Dominance Related to the Variability and Symmetry of the Vertical Ground Reaction Force and Center of Pressure. Journal of Applied Biomechanics. 2012;28(4):473-478. doi:10.1123/jab.28.4.473.

68. DeVita P, Hortobágyi T. Obesity is not associated with increased knee joint torque and power during level walking. Journal of Biomechanics. 2003;36(9):1355-1362. doi:10.1016/s0021-9290(03)00119-2.

69. Morbidity and Mortality Weekly Report (MMWR) | MMWR. Cdcgov. 2017. Available at: http://www.cdc.gov/MMWR. Accessed February 10, 2017.

70. Lenka P, Tiberwala DN. "Effect of stump length on postural steadiness during quiet stance in unilateral trans-tibial amputee." Al Ameen J Med Sci 3 (2010): 50-57. Available from: CINAHL Plus with Full Text, Ipswich, MA. Accessed December 22, 2016.

TABLES 39 39

Table 1: ICC Table Day I Day II

AB 1 595.7793 524.3557

AB 2 575.6115 581.8487

AB 3 573.6692 561.7358

AB 4 591.4796 571.7363

AB 5 509.2769 544.2491

ML 1 400.2551 381.976

ML 2 394.0069 393.7137

ML 3 357.859 341.1507

ML 4 369.2304 391.4755

ML 5 392.1136 390.3556

KM 1 403.2496 409.4607

KM 2 431.3596 397.3423

KM 3 427.6363 411.8651

KM 4 433.9063 422.2829

KM 5 429.3209 409.2053

MM 1 687.8264 704.0277

MM 2 695.2913 690.6416

MM 3 643.8624 649.4208

MM 4 655.0208 689.7019

MM 5 679.7253 693.0914

ICC 0.96

40 40 Table 2: Subject Demographics Intact Residual Limb Limb Step Time Since Height Weight Length Length Length Initial Age Sex (cm) (kg) BMI (cm) (cm) (cm) Amputation

51.2± F=1 94.66±1 94.66±1 30.04 92.5±13.5 68.54±12. 72.5±6.8 16.6±9.45

14.01 M=4 5.29 5.29 ±6.1 9 16 8

Table 3: MVIC During Manual Muscle Testing

Subject Intact GMax Prosthetic GMax Intact GMed Prosthetic GMed

1 673.01 308.873 1082.2 777.257

2 1143.48 268.652 856.032 360.554

3 3558.76 582.289 417.042 6263.68

4 419.949 113.597 209.914 96.9942

5 710.043 781.21 796.2 1430.77

Table 4: Average Muscle Activity During Gait Trials

Subject Intact GMax Prosthetic GMax Intact GMed Prosthetic GMed

1 6.0199 5.589 11.671 10.607

2 2.708 8.3 19.012 22.292

3 16.286 6.171 13.531 29.148

4 10.607 6.02 5.589 11.671

5 17.707 15.559 19.917 19.091

41 41 Table 5: Percent MVIC During Gait Trials

Subject Intact GMax Prosthetic GMax Intact GMed Prosthetic GMed

1 0.0195 0.018 0.011 0.014

2 0.002 0.031 0.022 0.062

3 0.005 0.011 0.032 0.005

4 0.025 0.053 0.027 0.12

5 0.025 0.02 0.025 0.013

Table 6: Average Forces During Heelstrike Prosthetic Subject Intact FX Prosthetic FX Intact FY Prosthetic FY Intact FZ FZ

1 194.404 95.874 92.402 71.974 980.965 899.489

2 175.476 124.093 98.084 73.381 1217.781 1131.731

3 197.347 128.987 69.548 65.196 1056.228 691.776

4 191.678 152.804 89.473 111.105 1261.461 1051.832

5 174.446 150.007 75.0 69.718 980.452 1011.218

Table 7: Average Forces During Midstance Prosthetic Subject Intact FX Prosthetic FX Intact FY Prosthetic FY Intact FZ FZ

1 18.378 15.084 62.74 54.324 590.023 584.558

2 14.316 12.069 61.101 55.273 3513.099 860.058

3 20.862 16.37 30.239 36.005 480.268 411.27

4 23.681 30.019 46.693 70.324 759.864 732.483

5 14.039 12.096 42.573 33.366 589.956 546.969

42 42

Table 8: Average Forces During Toe-Off Prosthetic Subject Intact FX Prosthetic FX Intact FY Prosthetic FY Intact FZ FZ

1 226.166 140.968 72.533 74.249 1022.71 888.775

2 228.247 176.769 112.989 90.127 1150.59 1121.032

3 248.73 133.627 42.881 53.061 950.193 459.123

4 254.35 165.637 86.569 58.505 1115.048 1112.305

5 143.585 152.96 67.857 57.765 832.182 848.866

Table 9: PCC Intact Limb

Intact FX AMP FX Intact FY AMP FY Intact FZ AMP FZ

Heelstrike

I GMax -0.011 0.357 -0.045 0.511 -0.187 0.678

I GMed 0.025 0.726 -0.648 0.099 0.305 -0.273

Midstance

I GMax 0.2 0.438 0.029 0.21 -0.605 0.711

I GMed 0.281 0.245 -0.867 -0.305 -0.095 -0.268

Toe-Off

I GMax -0.275 -0.032 -0.097 -0.431 -0.314 0.79

I GMed 0.124 -0.02 -0.339 -0.587 -0.199 -0.407

43 43

Table 10: PCC Prosthetic Limb

Intact FX AMP FX Intact FY AMP FY Intact FZ AMP FZ

Heelstrike

A GMax -0.099 0.493 0.533 0.949 0.832 0.678

A GMed -0.046 0.467 0.535 0.931 0.904 0.627

Midstance

A GMax 0.406 0.777 0.243 0.851 0.222 0.711

A GMed 0.450 0.783 0.21 0.863 0.287 0.721

Toe-Off

A GMax 0.479 0.73 0.619 0.097 0.639 0.79

A GMed 0.582 0.733 0.621 0.125 0.716 0.747

FIGURES 45 45

Figure 1: Prone Gluteus Maximus Testing (knee 90°)

Figure 2: Prone Gluteus Maximus Testing (knee 0°)

46 46

Figure 3: Sidelying Gluteus Medius Testing (knee 0°)

Figure 4: Sidelying Gluteus Medius Testing (knee 90°)

47 47

Figure 5: MVIC

Figure 6: Average Muscle Activity During Gait

48 48

Figure 7: Average GRF Heelstrike

Figure 8: Average GRF Midstance 49 49

Figure 9: Average GRF Toe-Off

APPENDIX: K-LEVELS 51 51

K-Level Functional Description Foot Description

K0 Does not have ability to ambulate or transfer Not eligible for prosthesis.

safely with or without assistance, and prosthesis

does not enhance quality of life or mobility.

K1 Has ability or potential to use prosthesis for External keel, SACH foot, or

transfers or ambulation on level surfaces at fixed single-axis ankle/foot.

cadence. Typical of limited and unlimited

household ambulatory.

K2 Has ability or potential for ambulation with ability Flexible-keel foot and multi-axial

to traverse low-level environmental barriers such ankle/foot.

as curbs, stairs, or uneven surfaces. Typical of

limited community ambulatory.

K3 Has ability or potential for ambulation with variable Flex-foot system, energy storing

cadence. Typical of community ambulatory who foot, multi-axial ankle/foot,

has ability to traverse most environmental barriers dynamic response, or flex-walk

and may have vocational, therapeutic, or exercise system or equal.

activity demands prosthetic use beyond simple

locomotion.

K4 Has ability or potential for prosthetic ambulation Any ankle/foot system

that exceeds basic ambulation skills, exhibiting appropriate.

high impact, stress, or energy levels. Typical of

prosthetic demands of child, active adult, or

athlete.