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DEVELOPMENT OF A COMPUTATIONAL MODEL OF

KNEE–ANKLE–FOOT COMPLEX FOR FOOT SUPPORT

DESIGN

LIU XUAN

Ph.D

The Hong Kong Polytechnic University

2013

lbsysText BoxThis thesis in electronic version is provided to the Library by the author. In the case where its contents is different from the printed version, the printed version shall prevail.

The Hong Kong Polytechnic University

Interdisciplinary Division of Biomedical Engineering

DEVELOPMENT OF A COMPUTATIONAL MODEL OF

KNEE–ANKLE–FOOT COMPLEX FOR FOOT SUPPORT DESIGN

LIU XUAN

A thesis submitted in partial fulfilment of the requirements for

the Degree of Doctor of Philosophy

November 2012

i

CERTIFICATE OF ORIGINALITY

I hereby declare that this thesis is my own work and that, to the best of my knowledge

and belief, it reproduces no material previously published or written, nor material that

has been accepted for the award of any other degree or diploma, except where due

acknowledge has been made in the text.

______________________

(Liu Xuan)

ii

ABSTRACT

Knee pain and functional impairment are the most common complaints among patients

with knee osteoarthritis (OA). Although the causative mechanisms of knee OA are not

entirely legible, there are evidences revealing that the initiation and progression of

degenerative processes at the knee are associated with joint loadings. The nature of

the knee loading is supposed to be altered by a number of conservative intervention

strategies, such as foot orthoses. The effects of foot orthoses on knee joint loading rely

mainly on experimental measurements. However, due to the experimental design

diversity, subject individual differences, and relatively small changes introduced by the

orthoses, consistent results have not been achieved. Furthermore, the knee adduction

moment (KAdM) was currently employed as a golden index of the knee loading

assessment in gait analysis, while leaving the loading distribution pattern inside the

joint unknown. Computer modeling, particularly finite element (FE) method gradually

manifests its advantage in exploring the biomechanical responses of joint internal

structures. Stress/strain distributions on the articulation surfaces predicted by FE

modeling could be considered as a more direct index of the knee loading. Thus,

whether the foot support reduces relevant knee compartment loading deserves further

deliberation through new perspectives.

For investigating the joint biomechanics and orthotic performance, an assessment

platform was established in this study, including gait analysis, musculoskeletal

modeling and FE modeling. Laterally wedged insoles (LWIs) with wedge angles of 0°,

5°and 10°were fabricated for orthotic interventions. Gait analyses were performed to

obtain necessary information to drive the musculoskeletal model and to setup FE

model. Musculoskeletal modeling was applied to calculate the muscle forces in each

LWI condition as FE loading boundary because the external loadings acting on joints

iii

must be balanced by muscle forces. To predict the stress, strain, pressure and force in

the bony and soft structures, a knee–ankle–foot FE model was developed. The model

consisted of 30 bony segments, including the distal segment of the femur, patella, tibia,

fibula, and 26 foot bones. The knee joint soft structures included the menisci, articular

cartilages, and the main ligaments of the knee joint. The established FE model was

partially validated through in-vivo plantar pressure and cadaveric tibiofemoral pressure

measurements.

From gait analysis, the KAdM peak was reduced by 16.1% and 19.6% in 5°and 10°

LWI conditions comparing with the 0°LWI condition, respectively. The decrease in the

KAdM during walking was directly related to a decrease of the medial compartment

loading at the knee joint. Actually, 5°and 10°LWIs relieved the KAdM prominently

during most of the stance phase, which demonstrated the effectiveness of the LWI

intervention on redistributing the knee joint loading from the experimentally measured

moment level. The musculoskeletal modeling results indicated that the gastrocnemius

lateralis force increased while gastrocnemius medialis force decreased in 5°and 10°

LWI conditions comparing with the 0°LWI condition, which implicated a strengthening

of the lateral muscle group spanning the knee joint triggered by LWI intervention.

In FE predictions, with either 5°or 10°LWI, there were significant decreases in stress,

strain, contact pressure and force at the medial femur cartilage and the medial

meniscus comparing with 0 ° LWI condition, which further demonstrated the

effectiveness of LWIs on redistributing knee joint loadings together with the gait

analysis results. The predicted maximum loadings on tibiofemoral articulation surface

during stance phase appeared at the second GRF peak, which was agreed with the

subject specific kinetic data. FE predictions also showed that the 5°and 10°LWIs

iv

reduced the lateral collateral ligament (LCL) force comparing with 0°LWIs. The

decreasing of the LCL force may be attributed to the increased muscle force spanning

lateral side of the knee joint induced by the LWIs. It was suggested that the

neuromuscular control made significant contributions to joint loading distributions and

LWIs interventions may have positive effects on prevention of the medial knee OA.

To our knowledge, there has been no FE model that attempts to incorporate the knee

and ankle–foot together considering the authentic motion features and muscle loadings.

In this study, FE modeling together with experimental studies and musculoskeletal

modeling successfully established a useful platform for understanding complicated joint

behaviors under different foot supports. Our experimental results and model predictions

also provided scientific fundamentals for evaluating and designing effective foot

supports.

v

PUBLICATIONS ARISING FROM THE THESIS Peer-reviewed Journals

1. Liu X, Zhang M, 2013. Redistribution of knee stress using laterally wedged insole intervention: finite element analysis of knee–ankle–foot complex. Clinical Biomechanics, 28: 61-67.

2. Liu X, Ouyang J, Fan YB, Zhang M, 2013. A foot–ankle–knee computational platform to explore foot support effects on knee joint biomechanics during gait. Annals of Biomedical Engineering, under review.

3. Liu X, Yu T, Ye JD, Zhang M, 2013. Investigation of mechanical behaviors of CPC/bone specimens by experiments combined with finite element analysis. Journal of American Ceramic Society, under review.

Conference Proceedings 1. Zhang M, Liu X, Yu J, Cong Y. Lower-limb biomechanics for foot support design. In:

Journal of Medical Biomechanics. Proceeding of 10th National Conference on Biomechanics and 12th National Conference on Biorheology, Chengdu, Sichuan, China, Oct.11-15, 2012, s27, pp8.

2. Liu X, Cheung JT, Zhang M. Influence of wedge insoles upon knee joint loading: using a finite element model of lower limb. In: Proceedings of the 5th World Association for Chinese Biomedical Engineers (WACBE) Congress on Bioengineering, Tainan, Taiwan, Aug 18-21, 2011.

3. Liu X, Zhang M. Finite element modeling of lower limb for impact attenuation. In: Proceedings of the 6th World Congress on Biomechanics, Singapore, Aug 1-6, 2010.

Book Chapters 1. Zhang M, Liu X. Ch. 13. Foot-ankle-knee model for foot orthosis. In: Computational

Biomechanics of Musculoskeletal System – Applications for Orthopaedics and Rehabilitation, Ed. Zhang M, Fan YB, CRC Press, 2013, in preparation.

vi

ACADEMIC AWARDS

2012 Student Research Award

Hong Kong Medical and Healthcare Device Industries Association

(HKMHDIA)

2008 Student Academic Competition Champion Award at

National Biomechanics Forum, in Shanghai by

The National Natural Science Foundation of China (NSFC)

vii

ACKNOWLEDGEMENTS

“To learn and to apply, for the benefit of mankind”

——PolyU Motto

This thesis is the milestone in the end of my journey for pursuing PhD study, but the

point of departure of my lifelong dedication in the world of biomedical engineering. The

most crucial thing PhD training teaches me was that talent is not everything, while

passion is something that will follow us as we put in the hard work to become valuable

to the world. I would never forget the Motto of PolyU, learn and apply, try the best to

improve human health employing my experience and imagination.

Day back from my first coming to PolyU four years ago, I still remember the scenario

that I had the first meeting with my supervisor Prof. Ming Zhang. His expectations and

erudite insights in biomechanics have continually inspired me through these years of

study. I would like to appreciate Prof. Ming Zhang very much for his profound kindness

and enormous support. I must also thank Prof. Yubo Fan, Dean of School of Biological

Science and Medical Engineering, Beihang University, China, who lighted my way in

transition from a mechanical engineer to a biomedical engineer at the very beginning of

my graduate study.

I would like to deliver special thanks to Prof. Jun Ouyang, Director of Clinical Anatomy

Institute, Southern Medical University, China for providing technical support on the

cadaveric experiments in this study.

viii

It is my pleasure to acknowledge all the current and previous staff and students in BME,

PolyU for giving me a warm and supporting environment throughout my PhD studies. I

would like to acknowledge our team members Dr. Jia Yu, Dr. Yan Cong, Dr. Tao Yu,

Ms. Yan Wang, Mr. Zhongjun Mo and Mr. Daiquan Zhang for their encouragement and

help in solving problems in my studies. I consider myself quite fortunate to be a

member of this research team. We enjoy a free atmosphere for discussing academic

topics and sharing ideas with each other, which have offered me an opportunity to

make progress subtly every day. I wish to thank Dr. Aaron Leung, Dr. Jason Cheung,

Dr. Zhihui Pang, Mr. Duo Wong, Ms. Meng Huang, Dr. Wenxin Niu, Dr. He Gong, Mr.

Yuxing Wang, Dr. Lizhen Wang and Dr. Yan Luximon, for their constructive comments

and recommendations given to me.

I also express my gratitude to the financial supports from the Research Grant Council

of Hong Kong (Project nos. PolyU 533107E, PolyU 535208E, PolyU 532611E) and

National Nature Science Foundation of China (11120101001, 11272273) for the grants

provided to this study.

Last but not least, I would like to pay high regards to my parents, grandparents, uncles

and aunts for their inspiration throughout my learning road and lifting me uphill this

phase of life. I love them very much. They are excellent people with wisdom, courage

and diligence, who give me the force of examples any time in my life. I owe everything

to them and I hope my family will be proud of me. Life is full of challenge and surprise

and I will keep walking.

ix

TABLE OF CONTENTS

CHAPTER I INTRODUCTION ....................................................... 1 1.1 Knee Osteoarthritis - A Public Concern .......................................... 1

1.1.1 What is Osteoarthritis? ........................................................................... 1

1.1.2 Prevalence of Osteoarthritis ................................................................... 5

1.1.3 Prevention and Treatments of Knee Osteoarthritis ................................. 7

1.1.4 Functions of Foot Orthoses .................................................................. 11

1.2 Objective of this Study .................................................................... 14 1.3 Outline of the Dissertation .............................................................. 16

CHAPTER II LITERATURE REVIEW .......................................... 18 2.1 Review on Human Lower Extremity Biomechanics ..................... 18

2.1.1 Functional Knee Anatomy .................................................................... 19

2.1.2 Functional Ankle and Foot Anatomy .................................................... 24

2.1.3 Biomechanical factors to knee joint degenerations .............................. 30

2.1.4 Biomechanics during gait ..................................................................... 33

2.1.5 Orthotic Intervention Evaluation ........................................................... 39

2.2 Review on Finite Element Analysis of Human Lower Extremity . 48 2.2.1 Introduction to Finite Element Analysis ................................................ 48

2.2.2 Principle and Procedure of Finite Element Analysis ............................. 49

2.2.3 History of Finite Element Analysis on Lower Extremity Research ........ 55

2.2.4 Summary of Existing Finite Element Analysis ...................................... 66

CHAPTER III METHODS ............................................................ 68 3.1 Gait Analysis .................................................................................... 68 3.2 Musculoskeletal Modeling .............................................................. 72 3.3 Development of the Finite Element Model .................................... 76

3.3.1 Geometry and Mesh ............................................................................. 77

3.3.2 Material Properties ............................................................................... 87

3.3.3 Loading and Boundary Conditions ....................................................... 96

3.4 Validation of the Finite Element Model ........................................ 100 3.4.1 Plantar Pressure Measurements ........................................................ 100

3.4.2 Cadaveric Experiments ...................................................................... 101

x

3.5 Summary ......................................................................................... 107

CHAPTER IV RESULTS ........................................................... 109 4.1 Gait Analysis: Kinematics and Kinetics ...................................... 109 4.2 Musculoskeletal Modeling: Muscle Forces ................................. 116 4.3 Validation of the FE Model of Knee–ankle–foot .......................... 123

4.3.1 Comparative Study with In-vivo Plantar Pressure Measurement ........ 123

4.3.2 Comparative Study with Cadaveric Experiments ................................ 130

4.4 FE Predictions for the Evaluation of LWI Intervention ............... 132 4.4.1 Effects on Tibiofemoral Articulation Loadings .................................... 132

4.4.2 Effects on Knee Ligaments ................................................................ 144

4.4.3 Effects on Foot and Ankle .................................................................. 145

CHAPTER V DISCUSSION ....................................................... 149 5.1 Gait Analysis: Kinematics and Kinetics ...................................... 149 5.2 Musculoskeletal Modeling: Muscle Forces ................................. 154 5.3 Validation of the FE Model of Knee–ankle–foot .......................... 158 5.4 FE Predictions for the Evaluation of LWI Intervention ............... 160

5.4.1 Effects on Tibiofemoral Articulation Loadings .................................... 160

5.4.2 Effects on Knee Ligaments ................................................................ 163

5.5 Limitations ...................................................................................... 165

CHAPTER VI CONCLUSIONS AND FUTURE WORK ............. 168 6.1 Conclusions ................................................................................... 168 6.2 Directions of Further Studies ....................................................... 171

REFERENCES. ............................................................................. 175

xi

LIST OF FIGURES

Figure 1-1. Knee Osteoarthritis .................................................................... 2 Figure 1-2. Radiographs of OA knees showing parallel MTP alignment.. .... 3 Figure 1-3. Joint symptoms and radiographic features of OA. ..................... 4 Figure 1-4. Toxicity profile of the treatment modalities based on expert

opinion. .................................................................................................. 8 Figure 1-5. The laterally wedged insoles ..................................................... 13 Figure 1-6. The KAdM during walking .......................................................... 14 Figure 2-1. Front view and back view of the main bony and soft structures

around the tibiofemoral joint .................................................................. 19 Figure 2-2. Front view and side view of the patella, quadriceps tendon (QT)

and patellar tendon (PT). ....................................................................... 23 Figure 2-3. Front view and back view of the main muscles around the knee

joint. ....................................................................................................... 24 Figure 2-4. Anatomy of the ankle and foot ................................................... 25 Figure 2-5. The medial longitudinal arch and windlass mechanism ............. 26 Figure 2-6. Side view and front view of the main muscles of the right leg .... 28 Figure 2-7. Back view of the main muscles of the right leg (from superficial to

deep layer) ............................................................................................. 29 Figure 2-8. The Q angle is formed between a line joining the anterior superior

iliac spine (ASIS) and the centre of the patella, and a line joining the centre of the patella and the tibial tuberosity. ................................................... 31

Figure 2-9. Normal gait cycle with approximated event timings. .................. 34 Figure 2-10. Six degrees of freedom (DOF) of the knee joint, which include 3

rotational and 3 translational motions. ................................................... 35 Figure 2-11. Joint motion of the foot/ankle in three anatomical planes. ....... 36 Figure 2-12. Muscular sequence of lower extremity muscles during gait.

Numbers signify percent of time in gait at which the muscle activity begins and ends ................................................................................................ 38

Figure 2-13. Flow chart of FEA solution process ......................................... 50 Figure 2-14. Sketch map for ABAQUS step sequence. ............................... 54 Figure 3-1. Wedged insole materials ........................................................... 69 Figure 3-2. Footwear with drilled holes used in the experiments ................. 70 Figure 3-3. Ground reaction force (GRF) measurement using AMTI force

platform (Advanced Mechanical Technology, Inc., MA, USA) and motion capture using the Vicon gait analysis system. ....................................... 71

Figure 3-4. Motion data analysis using Visual3D (trial version, C-motion Inc., MD, USA). ............................................................................................. 72

xii

Figure 3-5. 3D muscle-driven simulation of walking. .................................... 73 Figure 3-6. Schematic of the Computed Muscle Control Algorithm Applied to

Gait ........................................................................................................ 75 Figure 3-7. Screenshots from OpenSim. ...................................................... 76 Figure 3-8. Acquisition and reconstruction of knee–ankle–foot geometry. ... 77 Figure 3-9. Detailed geometries of the knee joint. Anterosuperior view. ...... 78 Figure 3-10. Software interfaces for solid model generation: (a) surface models

in Mimics v14 (Materialise, Leuven, Belgium), (b) solid models in Rapidform XOR3 (INUS Technology, Inc., Seoul, Korea)...................... 79

Figure 3-11. FE meshes of the tibiofemoral joint.......................................... 80 Figure 3-12. Three-dimensional FE model of the human knee–ankle–foot

complex together with wedged supports................................................ 81 Figure 3-13. Attachment points of the ligamentous structures ..................... 83 Figure 3-14. Master slave contact algorithm. Anchor point (X0) and tangent

plane are computed for every slave node based on the computed normal vectors. Each slave node (e.g. node 5) is constrained not to penetrate its tangent planes. ...................................................................................... 84

Figure 3-15. The 5° LWIs investigated in this study. (a) The fabricated 5° LWIs used in gait analysis. (b) The 5° LWI developed in the FE model. ......... 86

Figure 3-16. The FE mesh of the foot supports: (a) foot support with 0° LWI, (b) foot support with 5° LWI, (c) foot support with 10° LWI. ........................ 87

Figure 3-17. Stress–strain curves for various knee ligaments ..................... 90 Figure 3-18. ABAQUS evaluations of hyperelastic fitting for knee

ligaments.. ............................................................................................. 92 Figure 3-19. Stress-strain curve for plantar foot tissue. ............................... 93 Figure 3-20. ABAQUS evaluations of hyperelastic fitting for foot soft tissue.

Line with cross: experimental data; line with square: first-order polynomial form; line with circle: second-order polynomial form. ............................. 94

Figure 3-21. Schematic uniaxial stress-strain curve for elastomeric foam in compression (Gibson, 2005). ................................................................. 95

Figure 3-22. Model input differences explaining the lateral-medial shift of the COP under-neath foot support and the GRF directions in frontal plane with insoles at three wedge angles.. ............................................................. 97

Figure 3-23. Loading and boundary conditions of the FE model .................. 98 Figure 3-24. Connection type AXIAL. ........................................................... 99 Figure 3-25. F-Scan plantar pressure analysis ............................................ 101 Figure 3-26. K-Scan Sensor 4000 (Tekscan Inc., Boston, USA).................. 102 Figure 3-27. Preparation of cadaveric knee–ankle–foot specimen .............. 103 Figure 3-28. Insertion of K-Scan sensors ..................................................... 103 Figure 3-29. Tendon suture ......................................................................... 104

xiii

Figure 3-30. Pure compression tests ........................................................... 105 Figure 3-31. Compression tests combined with muscle loading: (a) balance

standing simulation from front view showing the muscle loading method through dead weights, (b) balance standing simulation from back view presenting Achilles tendon loading ........................................................ 106

Figure 3-32. Flow chart of FE model input, validation and output. Arrows with gradient fill indicate input and output, arrows with no fill shows validation .............................................................................................................. 106

Figure 4-1. COP trajectories in three LWI conditions during stance phase .. 110 Figure 4-2. GRF and shank-ground angle waveform data in three LWI

conditions during stance phase.. ........................................................... 111 Figure 4-3. Ankle joint and knee joint angle waveform data in three LWI

conditions during stance phase. ............................................................ 113 Figure 4-4. Ankle joint and knee joint moment waveform data in three LWI

conditions during stance phase. The moments were divided by body mass for normalization. ................................................................................... 116

Figure 4-5. Muscle forces during stance phase in 0°, 5°, and 10° LWI conditions. (a) gastrocnemius lateralis force, (b) gastrocnemius medialis force, (c) vasti force, (d) soleus force, (e) biceps femoris, (f) medial hamstring, (g) toe flexors, (h) peroneus, (i) tibialis posterior, (j) tibialis anterior. ................................................................................................. 123

Figure 4-6. The plantar pressure distribution at the first peak of GRF obtained from F-Scan (left column) and FE simulation (right column): (a) 0° LWI condition, (b) 5° LWI condition and (c) 10° LWI condition. ..................... 125

Figure 4-7. The plantar pressure distribution at the GRF valley obtained from F-Scan (left column) and FE simulation (right column): (a) 0° LWI condition, (b) 5° LWI condition and (c) 10° LWI condition. ..................................... 127

Figure 4-8. The plantar pressure distribution at the second peak of GRF obtained from F-Scan (left column) and FE simulation (right column): (a) 0° LWI condition, (b) 5° LWI condition and (c) 10° LWI condition. ............. 129

Figure 4-9. The K-Scan measured (left column) and FE predicted (right column) tibiofemoral pressure distributions during (a) pure compression test, (b) standing simulation and (c) mid-stance simulation. ............................... 131

Figure 4-10. von Mises stress in femur cartilage with (a) 0° LWI, (c) 5° LWI and (e) 10° LWI and von Mises stress in meniscus with (b) 0° LWI, (d) 5° LWI and (f) 10° LWI at GRF valley, using muscle force transformed from Perry’s work (1992). .............................................................................. 134

Figure 4-11. Effects of insole wedge angle on (a) average von Mises stress and (b) total contact force in medial meniscus and lateral meniscus at GRF valley, using muscle force transformed from Perry’s work (1992). ........ 135

Figure 4-12. von Mises stress in femur cartilage with (a) 0° LWI, (c) 5° LWI and (e) 10° LWI and von Mises stress in meniscus with (b) 0° LWI, (d) 5° LWI and (f) 10° LWI at GRF first peak, using muscle force calculated from musculoskeletal modeling. ..................................................................... 136

xiv

Figure 4-13. von Mises stress in femur cartilage with (a) 0° LWI, (c) 5° LWI and (e) 10° LWI and von Mises stress in meniscus with (b) 0° LWI, (d) 5° LWI and (f) 10° LWI at GRF valley, using muscle force calculated from musculoskeletal modeling. ..................................................................... 137

Figure 4-14. von Mises stress in femur cartilage with (a) 0° LWI, (c) 5° LWI and (e) 10° LWI and von Mises stress in meniscus with (b) 0° LWI, (d) 5° LWI and (f) 10° LWI at GRF second peak, using muscle force calculated from musculoskeletal modeling. ............................................................ 138

Figure 4-15. Minimum (compressive) principal strain in femur cartilage with (a) 0° LWI, (c) 5° LWI and (e) 10° LWI and min principal strain in meniscus with (b) 0° LWI, (d) 5° LWI and (f) 10° LWI, using muscle force calculated from musculoskeletal modeling. ............................................................ 139

Figure 4-16. Contact pressure on articulation surface of femur cartilage with (a) 0° LWI, (c) 5° LWI and (e) 10° LWI and contact pressure on articulation surfaces of meniscus with (b) 0° LWI, (d) 5° LWI and (f) 10° LWI, using muscle force calculated from musculoskeletal modeling. ...................... 140

Figure 4-17. Lateral-medial shear stress on tibial plateau with (a) 0° LWI, (c) 5° LWI and (e) 10° LWI, using muscle force calculated from musculoskeletal modeling.. .............................................................................................. 141

Figure 4-18. Tibiofemoral loading responses during stance phase, using muscle force calculated from musculoskeletal modeling: (a) peak von Mises stress in femoral cartilage, (b) peak von Mises stress in menisci, (c) peak min principal strain in femoral cartilage, (d) peak min principal strain in menisci, (e) peak contact pressure on femoral cartilage, (f) peak contact pressure on menisci, (g) peak lateral-medial shear stress in tibial plateau. ............................. 142

Figure 4-19. The ligament forces at the GRF first peak, valley and second peak, using muscle force calculated from musculoskeletal modeling. (a) ACL force, (b) LCL force. ....................................................................... 145

Figure 4-20. FE predicted von Mises stresses in the foot bones at the end of the mid-stance period with 0°, 5° and 10° LWIs, separately, using muscle force calculated from musculoskeletal modeling.. ................................. 147

Figure 4-21. FE predicted peak von Mises stresses in the foot bones at the end of the mid-stance period with 0°, 5° and 10° LWIs, separately, using muscle force calculated from musculoskeletal modeling. ...................... 148

xv

LIST OF TABLES Table 1-1. Radiographic grades of severity for OA of the knee (Kellgren,

1963). .................................................................................................... 4 Table 1-2. Determinants of progression of hip and knee OA. (Arden and Nevitt,

2006). .................................................................................................... 5 Table 2-1. Summary of research about wedge foot orthotics effects on the

knee joint. .............................................................................................. 40 Table 2-2. Summary of FE models of knee joint structures in literatures ..... 57 Table 3-1. Material properties and element types defined in the FE model . 89 Table 3-2. The coefficients of the hyperelastic material model used for the

knee ligaments ...................................................................................... 91 Table 3-3. The coefficients of the hyperelastic material model used for the foot

encapsulated soft tissue ........................................................................ 93 Table 4-1. Peak muscle force and occurrence time for different muscle groups

from EMG data (Perry, 1992) and the calculation in current study based on the musculoskeletal model. ................................................................... 118

Table 4-2. FE predicted peak contact pressure and contact area in the medial compartment and lateral compartment of the knee compared with the mean peak contact pressure and contact area (SD) obtained from the cadaveric experiment (Poh et al., 2011) ................................................ 132

Table 4-3. Muscle forces for single-limb stance simulation in preliminary study .............................................................................................................. 133

xvi

LIST OF ABBREVIATIONS

ACL: anterior cruciate ligament

3D: three-dimensional

AP: anterior-posterior

BF: biceps femoris

BW: body weight

COP: center of pressure

EMG: electromyography

EVA: ethylvinylacetate

FDL: flexor digitorum longus

FE: finite element

FEA: finite element analysis

FHL: flexor hallucis longus

GAS: gastrocnemius

GL: gastrocnemius lateralis

GM: gastrocnemius medialis

GRF: ground reaction force

HAM: hamstring

KAdM: knee adduction moment

LCL: lateral collateral ligament

LWI(s): laterally wedged insole(s)

MCL: medial collateral ligament

MH: medial hamstring

ML: medial-lateral

MR: magnetic resonance

OA: osteoarthritis

PB: peroneus brevis

PER: peroneus

PCL: posterior cruciate ligament

PL: peroneus longus

QUADS: quadriceps

SOL: soleus

TA: tibialis anterior

TOEF: toe flexor

TP: tibialis posterior

VAS: vasti

Chapter I Introduction

1

CHAPTER I INTRODUCTION

1.1 Knee Osteoarthritis - A Public Concern

1.1.1 What is Osteoarthritis?

Osteoarthritis (OA), also known as degenerative arthritis or degenerative joint disease

or osteoarthrosis, could be defined as a heterogeneous group of conditions that lead to

joint symptoms and signs which are associated with defective integrity of articular

cartilage, in addition to related changes in the entire joint. OA can happen in almost

every joint; however, knee OA is the most common musculoskeletal complaint that

brings people to their doctors. Knee OA, as a metabolically active, dynamic disease

includes both destruction and repair mechanisms that may be triggered by biochemical

and mechanical insults (Astephen et al., 2008). Since the knee joint plays a main role

in weight bearing activities, it is more susceptible to mechanical insults than other joints.

The mechanical abnormalities of knee OA manifest degradation of joints, mainly

including articular cartilages and subchondral bones, as shown in Fig. 1-1. The

surfaces of the subchondral bone are covered by cartilage. The surfaces allow the

bones to slide against each other without causing damage to the bone. Thus, one of

the main roles of the articular cartilages is to protect the subchondral bone from impact

and attrition. When the cartilage is degraded, the subchondral bone will expose to the

external stimuli.

http://en.wikipedia.org/wiki/Joint�http://en.wikipedia.org/wiki/Articular_cartilage�http://en.wikipedia.org/wiki/Subchondral_bone�http://en.wikipedia.org/wiki/Articular_cartilage�http://en.wikipedia.org/wiki/Subchondral_bone�


2

Figure 1-1. Knee Osteoarthritis (Osteoarthritisblog.com, 2012)

OA may occur as a secondary effect of traumatic joint injury or other abnormal joint

loading condition, or it may arise from normal repetitive movement (also known as

“wear and tear”) on joints with no other primary cause. Either way, with OA there is

erosion of the cartilage, which could give rise to a succession of joint environment

alteration. Clinically, the symptoms are characterized by joint pain, tenderness,

limitation of movement, crepitus, occasional effusion, and variable degrees of local

inflammation.

The initial changes in OA occur at the microscopic level with changes in collagen and

macromolecular components of the extracellular matrix. Histologically, the disease is

characterized early by fragmentation of the cartilage surface, cloning of chondrocytes,

vertical clefts in the cartilage, variable crystal deposition, abnormal remodeling, and

eventual violation of the tidemark by blood vessels (Brandt et al., 1986). These

changes compromise the cartilage mechanical properties, leading to further disruption

until the characteristic symptoms of pain and stiffness are experienced (Buckwalter et

al., 2008).


3

All tissues of the joint are involved along with the onset and progression of OA, to a

certain extent with ligamentous laxity and weakening of muscles around the joint.

Actually, soft tissue stress reduction could adversely affect the ability to control joint

movements precisely. In this regard, OA represents failure of the joint as an integral

organ more than as the bony tissue only (Arden and Nevitt, 2006). The most widely

using classification schemes for OA are in accordance with the radiological appearance

of the joint. The radiographic features conventionally used to define OA include joint

space narrowing, osteophytosis, subchondral sclerosis, cyst formation, and

abnormalities of bone contour. Nevertheless, radiographs could not show other

important structures that account for pain in OA. Symptomatic knee OA is more

common within the medial tibiofemoral compartment than within the lateral

compartment. Figure 1-2 depicts the alignment of the medial tibial plateau (MTP) and

joint space narrowing in the OA knee.

Figure 1-2. Radiographs of OA knees showing parallel MTP alignment (Vignon et al., 2010). (a) with superimposition of the anterior and posterior margins (arrows) of the plateau (IMD = 0 mm at the middle of the MTP); (b) skewed alignment of the MTP (IMD = 2.5 mm in the middle of the medial joint space) and (c) dramatic misalignment (IMD = 7.5 mm in the middle of the medial joint space) *IMD means inter-margin distance.


4

Table 1-1 displays the categories for knee OA developed by Kellgren (1963). Based

solely on radiographic differences comparing with normal atlas, grades are classified

as zero-absent, 1-doubtful, 2-minimal, 3-moderate and 4-severe. Besides the

radiographic comparison, symptomatic judgement and self-reported assessment are

also commonly used for the OA diagnosis. Joint pain is the dominant symptom of OA.

The association between joint pain and radiographic features of OA is not constant. In

studies performed during the 1950s in the north of England (Fig.1-3), the relationship

between pain and radiographic evidence of OA was considerably stronger for the hip

and knee than for distal interphalangeal (DIP) joint involvement (Lawrence, 1977).

Table 1-1. Radiographic grades of severity for OA of the knee (Kellgren, 1963).

Figure 1-3. Joint symptoms and radiographic features of OA. (Lawrence, 1977)


5

OA is most commonly found in the knee, hip, spine and hand. Wrist, elbow, shoulder

and ankle can also be affected by OA, but occur less frequently. Just as the generation

history of OA differs at different joint sites, the factors that contribute to disease

progression also appear to be joint specific. Table 1-2 summarizes the known

determinants of progression at the knee and hip. At both sites, multiple joint

involvements appear to be a determinant of accelerated disease. It is worth noting that

though the risk factors listed here are critical in some degree, the authentic nature by

which they contribute to joint degenerative changes remain speculative. For some risk

factors such as mal-alignment, it exacerbated the symptoms of OA and in the

meantime intensified by OA progression. Whether it represents an independent risk

factor for disease onset and rate of progression is not clear. Thus, further investigations

on the biomechanical disease mechanism should be considered.

Table 1-2. Determinants of progression of hip and knee OA. (Arden and Nevitt, 2006).

1.1.2 Prevalence of Osteoarthritis

Osteoarthritis is among the most prevalent of human musculoskeletal diseases

(Pereira et al., 2011). It imposes a significant economic burden and is associated with

pain, disability and loss of quality of life. The prevalence of OA increases with age and

there are differences between genders. Males are affected more frequently than

females below age 45, while females are affected more often after age 55 (Silman and

Hochberg, 1993). OA increases dramatically with years affecting nearly 27 million


6

Americans (Helmick et al., 2008) and 151 million individuals worldwide (Symmons et al.,

2006).

Most currently available information on the epidemiology of OA come from population

based radiographic surveys. Although the majority of studies have focused on the

prevalence of radiographic changes in OA, there is an increasing amount of research

into the prevalence of self-reported joint pain. Prevalence estimates from pathological

studies tend to be higher than those from radiographic surveys, partly because

relatively mild pathological change is not apparent on radiographs, and because

pathological studies examine the complete joint surface.

Most of the OA disability burden is attributable to the hip and knee, since pain and

stiffness in large weight bearing joints often lead to significant problems. In fact, OA is

the precipitating diagnosis for more than 90% of the increasing number of total hip or

knee joint replacement operations being undertaken worldwide (Australian Orthopaedic

Association, 2009). In the US alone, the combined number of hip and knee joint

replacements performed is in excess of 350,000 annually (Arden and Nevitt, 2006).

The incidence rate in Asian countries is as high as western countries. Zhang et al

(2001) compared the prevalence of knee OA in elderly population between Beijing,

China and Framingham, US and found that women in Beijing had a higher prevalence

of radiographic knee OA and symptomatic knee OA. Several recent large population-

based cohort studies conducted in Asian countries have further confirmed an increased

risk of symptomatic knee OA associated with ageing, obesity and unhealthy lifestyle

(Fransen et al., 2011).

In 2007, OA increased the probability of missed workdays by 14% in women and by

12% in men. The effect of OA on workday loss was larger than that other common


7

conditions such as anxiety disorder, asthma, or diabetes (Kotlarz et al., 2010). As

incidence and prevalence rise with ageing, extending life expectancy will result in

greater numbers with OA.

1.1.3 Prevention and Treatments of Knee Osteoarthritis

Knee Osteoarthritis (OA) is the most prevalent OA with progressive and complicated

symptoms that could further lead to disability of the lower extremity. A World Health

Organisation report (Murray and Lopez, 1997) indicated that knee OA is likely to

become the fourth most important global cause of disability in women and the eighth

most important in men. The annual costs attributable to knee OA are immense.

Therefore, there is a burden for prevention and treatments of knee OA.

Among the large amounts of risk factors, local biomechanical factors (overweight,

occupational factors, sports participation, muscle weakness, etc.) have to do either with

risk of trauma or repetitive high loading on the knee joint and provide entry points for

prevention of knee OA. Other factors (for example, ageing, sex and heredity) are

systemic and could not be modified through prevention. Severe joint injury may be

sufficient to cause OA; however, the disease is often the product of interplay between

systemic and local factors. For example, a person may have an inherited predisposition

to develop the disease but will develop it only where a biomechanical insult has

occurred (Felson et al., 2000).

Treatments for knee OA come down to several choices. Surgical treatments comprise

joint replacement and osteotomy, while nonsurgical treatments consist of drug therapy,

intra-articular injection and physiotherapy. Figure 1-4 shows the toxicity profile of the

treatment modalities based on expert opinion (Jordan et al., 2003). From the histogram,

we could find that the toxicity of the surgical and pharmacological intervention is higher

than non-surgical and non-pharmacological intervention. The risks of some of the


8

interventions will be discussed in branch sections below.

Figure 1-4. Toxicity profile of the treatment modalities based on expert opinion. (Jordan et al., 2003)

Joint Replacement

The knee replacement operation is a surgical procedure aimed at replacing the worn

out parts of the arthritic knee with specially designed metal and plastic components. It

can be carried out as a total knee replacement or as a partial knee replacement. In the

most serious OA stage, knee replacement operation is sometimes the only surgical

option available to patients. And it is widely adopted for elderly patients to relieve pain

and restore joint functions. However, such procedures have higher failure rates in

young and middle-aged patients. This is due to the higher physical activity level of

these patients, which results in wear and loosening of the prosthesis. There is

preliminary evidence that achieving optimal lower extremity muscle strength, gait

patterns and symmetry, and knee joint loading patterns may be crucial for long-term


9

clinical benefits, integrity of other lower extremity joints and implant longevity after knee

replacement surgery.

Osteotomy

An osteotomy is a type of surgery in which the bones are cut and realigned. Re-

alignment can be achieved by either taking a slice of bone out of the tibia or femur

close to the knee joint or opening a gap in the bone. Osteotomy is frequently used as

treatment for unicompartmental OA. High tibial osteotomy is a treatment option for

medial OA of the knee and the goal of this operation is to reduce the abnormal loading

on the medial compartment of the knee. It is thought that an osteotomy may decrease

pain, improve function, slow damage in the knee, and possibly delay the need for

partial or total knee replacement surgery. Suitability for osteotomy surgery depends on

a number of factors which include age, bone alignment, activity level and isolated

damage to one side of the knee. Where a patient is suitable an osteotomy may

represent an excellent alternative to knee replacement surgery.

Drug Therapy

Paracetamol is recommended as the first choice in drug therapy for OA patients with

mild to moderate pain. NSAIDs (Nonsteroidal anti-inflammatory drugs) are widely used

in knee OA patients with severe pain and in those who have failed to respond to

paracetamol. However, NSAIDs can cause stomach upset, cardiovascular problems,

bleeding problems, and liver and kidney damage. Besides, Schnitzer el al (1993) found

that NSAIDs reduced symptomatic pain but increased knee joint loading in OA patients.

The increased loading may be attributed to pain free activity.

Intra-articular Injection

In OA, a reduction in the viscosity of the synovial fluid, secondary to a decrease in the

molecular weight and concentration of hyaluronic acid occurs. As a result, the synovial


10

fluid has decreased lubricating properties. Intra-articular injection of exogenous high-

molecular-weight hyaluronic acid molecules is an effective technique to increase the

synovial viscosity (Yang et al., 2004). For intra-articular injection, doctors should pay

attention that the drug is injected in the correct location within the joint.

Physiotherapy

Physiotherapy is a non-pharmacological conservative treatment approach that is

recommended in clinical guidelines for the management of knee OA. Physiotherapy

that aims to achieve modification of joint loading or cartilage structure seems to be

more promising in at-risk individuals or those with early disease. Physiotherapy is non-

invasive, economical and with minimal side-effects and should be considered prior to

surgical intervention. If satisfactory relief can be achieved using this option, risks

associated with other treatments are avoided. Physiotherapy is also applied as a

prevention process because of the widespread convenience it can offer most patients

in their daily life. Actually, prevention is the most effective treatment strategy and

patients with knee OA could try to monitor and control their symptoms through

physiotherapy. The most widely used physiotherapy comprise exercise, manual

therapy, braces and knee taping, orthotics and footwear, gait retraining and

electrotherapy. Among these treatments, orthotic intervention efforts on adjusting knee

loading through foot support alteration have been investigated for decades. However,

due to the relatively small changes introduced by the orthoses and subject individual

differences, consistent results on the orthotics effects around the proximal knee joint

are not yet to be achieved (Abdallah and Radwan, 2011; Kakihana et al; 2005;

Kerrigan et al., 2002; Kakihana et al., 2007; Maly et al., 2002). Since footwear and

orthotics play an important role in lower extremity care, it is interesting to understand

the mechanism through what they influence the knee behavior and how much they will

change the knee loading by foot support parameter alterations.

http://www.ncbi.nlm.nih.gov/pubmed?term=%22Maly%20MR%22%5BAuthor%5D�


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1.1.4 Functions of Foot Orthoses

Orthosis is an externally applied device used to modify the structural or functional

characteristics of the musculoskeletal system. Foot orthotics, as an important part of

podiatry and orthopedics non-pharmacological cure, have been widely used to improve

gait mechanics, heal injury to the foot, leg, and lower back. Since Whitman made a

metal foot brace which was proposed to push the foot into proper position (Whitman R,

1889), foot orthoses have been used for over a century by clinicians according to

literature records.

Functions of foot orthoses have been identified in ISO #8549-1 (1989): prevent, reduce,

or stabilize a deformity; modify the range of motion of a joint; add to the length or alter

the shape of a segment; compensate for weak muscle activity or control muscle

hyperactivity; and reduce or redistribute the load on tissues. The foot orthoses are

usually easy to be fabricated and constructed from a relatively inexpensive material.

In general, foot orthoses fall into one of two broad categories: functional or

accommodative. Functional orthoses seek to control the subtalar joint and foot

biomechanics, while accommodative orthoses minimize changes to foot function while

providing relief and protection to specific areas of the foot. Functional foot orthoses are

usually made from thinner, firmer materials, for example polypropylene. Usually they

will incorporate a deep heel cup and a good medial longitudinal arch. Among other

diagnoses, functional devices are used to treat pronation, plantar fasciitis, and heel

spur syndrome. Accommodative devices tend to be made from less rigid materials

such as Ethylene Vinyl Acetate (EVA). They are usually molded to the entire plantar

surface of the foot, providing comfort. Accommodative orthoses are a good choice for

patients with diabetes or a limited range of motion.

There is considerable research evidence that supports the therapeutic efficacy and

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12

significant mechanical effects of foot orthoses on daily activities such as standing,

walking and running. Although the most direct effects of the orthoses focus on

structural foot deformities, the use of foot orthoses should be considered as an adjunct

to treatment of lower extremity dysfunction related to poor alignment and faulty

mechanics (Nawoczenski and Epler, 1997). The fact that patterns of ambulation can be

modified by external stimuli has been used to study the influence of orthoses and

alteration of orthoses parameters on the loading at the knee.

There is great promise for increased understanding and further development of foot

orthoses as a valuable therapeutic tool in the treatment of mechanically based

musculoskeletal diseases (Matthew et al., 2010). The integral role of biomechanical

factors in the development and progression of knee OA is becoming widely

acknowledged. Surveys about therapeutic effects of foot orthoses on knee OA have

been conducted by clinicians for decades. However, theoretical explanations of how

orthoses actually produce their effects in mechanical way need to be continually refined.

Wedge orthoses are a category of foot orthoses which are prevalent in the

conservative treatment of knee OA. Wedge insole is a wedge inclined along a

particular side of the foot. Laterally wedged insole (LWI), with full length elevation on

the lateral side as shown in Fig. 1-5, has been suggested as an intervention strategy to

reduce the knee adduction moment (KAdM) during walking and attenuate the

progression of medial knee OA.


13

Figure 1-5. The laterally wedged insoles (Hinman el al., 2009).

Since knee OA is closely associated with the development of a high KAdM, which

reflects compression of the medial compartment of the knee, KAdM has been shown to

be a main predictor for medial knee OA. Throughout the stance phase of walking, a

KAdM tends to rotate the tibia medially with respect to the femur in the frontal plane

(Fig. 1-6a). This KAdM is primarily caused by a medially acting ground reaction force

(GRF). As shown in Fig. 1-6b, a varus knee deformity is strongly associated with a

higher KAdM. It alters the forces at the knee so that the line of force shifts farther

medially from the knee joint centre intensifying the already high medial compartment

loadings. Fig. 1-6c indicated that, when a LWI was applied, it could diminish the KAdM

by causing the GRF to pass closer to the knee joint centre. However, the efficacy, the

factors of the parametrical influence (inclination angles, etc.) of the orthoses are not

clear. Quantitative research on joint loading alterations induced by such a conservative

treatment should be launched. Evidences and parameters should be obtained through

a new perspective.


14

Figure 1-6. The KAdM during walking (Reeves and Bowling, 2011).The magnitude and direction of the GRF are shown by the height and direction, respectively, of the straight arrows. The length of the moment arm of the GRF acting about the knee joint is indicated by dotted red lines.

1.2 Objective of this Study

The biomechanical factors play an important role on the etiology, treatment and

prevention of many lower extremity disorders. It is essential to understand the

biomechanics associated with the normal lower extremity alignment before any foot

orthotic intervention can be applied. Information on the internal stress/strain of the knee

joint structures is essential in enhancing knowledge on the biomechanical behaviors of

the knee–ankle–foot complex. Direct measurement of those parameters is difficult,

while a comprehensive computational model can acquire those important clinical

information with more flexibility.


15

In recent literatures, many experimental techniques were developed and employed for

the quantification of lower extremity biomechanics, such as gait analysis systems,

pressure sensing platforms, in-shoe pressure transducers, pressure sensitive films, in-

vivo implant and cadaveric measurements. The above-mentioned measurement

techniques are commonly used in predicting joint kinetics and kinematics, and

quantifying pressure distributions on any internal or external contact surfaces. However,

stresses on cartilages, menisci, bones and associated structures inside the knee joint

are not well addressed. It is difficult to quantify the in-vivo bone and soft tissue stress

with the existing experimental techniques. As for in-vitro studies, the loading conditions

were often different from the actual physiological loading situations. Therefore, no

overall stress distribution of the whole knee joint during gait with foot orthotic

intervention is known using the currently available measuring techniques.

Apart from the experimental approaches, many theoretical models, such as kinematics

models, mathematical models, and finite element (FE) models of the knee have been

developed. FE method has been used increasingly in many biomechanical

investigations with great success due to its capability of modeling structures with

irregular geometry and complex material properties, and the ease of simulating

complicated boundary and loading conditions in both static and dynamic analyses.

Therefore, it has become a suitable method for the investigation of loading distribution

of the human joint. Although many FE analyses of the knee or foot were performed in

the literature, the knee–ankle–foot as a complex was seldom considered. Even with an

FE model of the complete lower extremity, only simplified geometry and loading

conditions were involved for the specific research interests. Therefore, a detailed FE

model is essential to provide an overall representation of the knee–ankle–foot and to

simulate the orthotics in gait.


16

Thus, the primary objective of this study is to establish a comprehensive FE model of

the human knee–ankle–foot to quantify the biomechanical responses of knee

structures under various wedge insole supports in gait. It is expected that the

developed FE model can provide quantitative information on stress/strain among the

bony and soft tissue structures, and pressure distribution on the articulation surfaces.

Besides the FE modeling, musculoskeletal modeling was also employed in this study.

The calculated muscle forces were used as input for the FE analysis. Considering the

muscle activation difference among orthoses conditions is supposed to be important

because the knee joint is vulnerable to intrinsic stability controlled by the active muscle

contractions and passive ligament tension.

Nowadays, evaluations of foot orthotic intervention for knee problems mainly rely on

clinical assessment (such as radiography and pain assessments) and KAdM

measurement through gait analysis. The biomechanical indices of the knee joint altered

by foot orthoses are still indistinct. In this study, the inclination levels of LWIs will be

investigated with an attempt to provide knowledge about the effects of wedge orthoses

on knee joint loading distributions and knee–ankle–foot biomechanical responses.

In summary, this study aimed to achieve the following objectives:

1. To develop a comprehensive FE model of the human knee–ankle–foot complex;

2. To establish a platform including both experimental and computational methods to

assess the effects of LWIs on knee joint loadings;

3. To build connections between foot support design and knee joint biomechanics.

1.3 Outline of the Dissertation

Following the introduction chapter, chapter II begins with a review of the functional

anatomy of the knee, ankle and foot. The biomechanical factors to knee joint


17

degenerations, biomechanics during gait and orthotic intervention evaluation are then

reviewed. Finally, the principle and advantages of FE method are introduced and

followed by a detailed review on the existing knee FE models.

In chapter III, the gait analysis and the musculoskeletal modeling procedures are first

prescribed. After that, the development of the FE model and the simulated conditions

are presented in details. The geometrical and material properties defined in the FE

model are delineated. The loading and boundary conditions applied for simulating the

physiological loading conditions of the human knee–ankle–foot complex are discussed.

Then, experiments conducted to validate the FE predictions are described.

Chapter IV presents the results of the gait analysis, the musculoskeletal modeling and

FE analysis. This chapter reports the findings of the LWI effects using FE predictions

together with gait analysis data and musculoskeletal calculations. The results of

cadaveric experiments and plantar pressure measurements for FE model validations

are included.

Chapter V discusses the effects of different LWI inclination angles on the

biomechanical responses of knee–ankle–foot structures. Relevant clinical implications

of the wedge orthoses are presented. In the last section of this chapter, the limitations

of this study are discussed.

Chapter VI summarises the findings in this study and its clinical implications regarding

different simulated conditions of the orthoses. Suggestions on further development of

the FE model and future research directions are highlighted.

Chapter II Literature Review

18

CHAPTER II LITERATURE REVIEW

2.1 Review on Human Lower Extremity Biomechanics

The Biomechanics research is using mechanical methods to solve the structural and

functional problems of biological systems. The understanding of biomechanics of the

lower extremity could provide researchers with invaluable information about the human

musculoskeletal system. Actually, lower extremities play significant roles in standing,

walking, jumping, running and similar activities, and constitute a vital portion of the

body weight (BW). Therefore, the disturbances of normal mechanical functions of the

lower extremity would lead to a wide variety of pathologic conditions, which

consequently cause inconvenience and suffering during sports and daily movements.

Besides improving the ambulatory function obviously, lower extremity biomechanics

continue to expand role in clinical issues. As an ever-increasing emphasis has been

placed on preventive and management of medical problems, accurately diagnosing

and conservatively treating of lower extremity problems have never been more

important. Therefore, biomechanics is also a reliable basis for the clinician to make a

decision when facing kinds of treatments or surgeries.

For better exploring how foot orthoses relieve the proximal lower extremity joint

problems such as knee pain, it is important to be familiar with the anatomy and function

of the knee. In this subchapter, we will first introduce the anatomy of the lower

extremity structures including knee and ankle-foot anatomy. With this foundation, we

will conduct review on the biomechanical factors to knee joint degenerations, classical

gait theories and up-to-date research on wedge orthotics intervention for knee

problems.

http://en.wikipedia.org/wiki/Standing�http://en.wikipedia.org/wiki/Walking�http://en.wikipedia.org/wiki/Jumping�http://en.wikipedia.org/wiki/Running�


19

2.1.1 Functional Knee Anatomy

The knee joint is the largest and perhaps most complex joint in the human body. It

comprises the tibiofemoral joint (Fig. 2-1) and the patellofemoral joint. The high

incidence of OA at the knee can be related to its intricate anatomy and has a strong

relationship with the mechanical loading. The normal function of the knee joint is

dependent on the interaction between the articular surfaces, the function of the

ligaments and menisci and the extrinsic forces (gravitational, inertial and muscle

generated) that act at the knee during ambulation (Andriacchi and Dyrby, 2005). The

individual contributions of internal structures are influenced by the anatomic features of

each structure and their capacity for absorbance and distribution of joint loading.

Figure 2-1. Front view and back view of the main bony and soft structures around the tibiofemoral joint (Interactive Series, 1999).

Femur

The femur is the longest bone in the body. Its proximal end comprises a head, neck,

greater trochanter and lesser trochanter. The lower end of the femur, which constitutes

the tibiofemoral articulation, comprises the lateral and medial femoral condyles. Each

condyle is curved from front to back forming shape like cam. The femoral condyles are

confluent anteriorly and separated from each other posteriorly by a deep notch termed

the intercondylar fossa.


20

The area of the femoral condyles is covered by articular hyaline cartilage. The cartilage

offers a firm and smooth surface to facilitate joint movements and protect the

subchondral bone. The surface of the cartilage is lubricated by synovial fluid secreted

by the synovial membrane. The thickness of cartilages in the knee is not uniform that

varies due to regional functions. The compressibility and elasticity of the cartilage

enable the dissipation of compressive forces and permit smooth glide for joint surface.

Tibia

The articular surface of the superior tibia is termed the tibial plateau. The medial and

lateral articular facets are covered by articular hyaline cartilages separately. The

surface is larger and more concave on the medial tibial plateau and this allows the

more curved medial femur condyle to rotate with ease. Anteriorly below the tibial

plateau, the tibial tuberosity is characterized by the presence of an irregular

prominence.

Fibula

The fibula is a long bone on the lateral side of the tibia. The proximal end of the fibula

is connected toward the back of the tibia head. The articulation surfaces are below the

plane of the tibiofemoral joint. The distal end of the fibula inclines anteriorly relative to

the proximal end and projects below the tibia end forming the lateral part of the ankle

joint.

Menisci

There are two menisci termed medial meniscus and lateral meniscus, which located

between the femur cartilage and tibia cartilages. They are both crescentic

fibrocartilaginous structures. The medial meniscus is approximately semi-circular and

wider posteriorly than anteriorly. The lateral meniscus is approximately circular and

http://en.wikipedia.org/wiki/Bone�http://en.wikipedia.org/wiki/Lateral_(anatomy)�http://en.wikipedia.org/wiki/Tibia�http://en.wikipedia.org/wiki/Knee�


21

with uniform width along the anterior-posterior (AP) direction.

In spite of the peripheral attachments, the menisci retain sufficient mobility to adapt to

the movements of the knee joint. The outer border of each meniscus is thicker than the

inner free border, forming a wedge-shape in its cross section. Each meniscus

possesses an anterior horn and a posterior horn. The horns of the menisci are attached

to the non-articular intercondylar area of the tibial plateau. The upper surface of each

meniscus is concave and articulates with the peripheral part of the overlying femoral

condyle. The lower surface of the meniscus is flat and locates on the articular facet of

the tibial plateau.

The meniscus plays significant role in bearing the tibiofemoral loading. A large portion

of the compressive loading of the joint is transmitted through the menisci no matter the

knee is in extension or flexion. Meniscal tears are among the most common traumatic

lesions of the knee joint. Medial meniscal tears are more commonly observed in stable

knee joints. While lateral meniscal tears are of a more common occurrence with

anterior cruciate ligament (ACL) injuries.

Cruciate Ligaments: ACL and PCL

The cruciate ligaments are named anterior and posterior with regard to the positions of

their attachments on the tibial plateau. They are called cruciate ligaments because they

cross each other like the letter X. Both ligaments are tense in the flexed position of the

knee. The ACL is attached to the anterior intercondylar area of the tibial plateau, with

the function to resist anterior translation and medial rotation of the tibia. The posterior

cruciate ligament (PCL) connects the posterior intercondylar area of the tibia to the

medial condyle of the femur. This configuration allows the PCL to resist posterior

translation of the tibia. Damage to the cruciate ligaments is particularly threatening to

knee stability.

http://en.wikipedia.org/wiki/Posterior_intercondylar_area�http://en.wikipedia.org/wiki/Tibia�http://en.wikipedia.org/wiki/Medial_condyle_of_femur�http://en.wikipedia.org/wiki/Femur�


22

Collateral Ligaments: MCL and LCL

The medial collateral ligament (MCL) is situated on the medial side of the knee joint

and composed of superficial and deep fibers. It connects medial condyle of femur

approximately below the adductor tubercle to the medial condyle of the tibia. The MCL

is fused with the joint capsule and the medial meniscus. It resists forces that push the

knee medially, which may lead to valgus deformity. The lateral collateral ligament (LCL)

is situated on the lateral side of the knee joint and was narrower than the MCL. It

connects lateral condyle of femur approximately below the lateral femur epicondyle to

the head of the fibula. The LCL is more flexible less susceptible to injury than the MCL

due to absence of attachment points on joint capsule or meniscus. The MCL and LCL,

as well as the anterior part of the ACL, are tense with the knee extended.

Patella, Quadriceps tendon (QT), Patellar tendon (PT)

The patella is the largest sesamoid bone in the body. It is located anterior to the

tibiofemoral articulation and protects the anterior articular surface of the knee joint. The

posterior surface of the patella with the femur and soft tissues between them form the

patellofemoral articulation, which is the second articulation of the knee joint besides the

tibiofemoral articulation. The patella is an important part of the knee joint extensor

mechanism. It lies embedded in QT that increases the arm of the quadriceps muscle

force. The PT is the inferior continuation of QT that attached to the lower apex of the

patella (Fig. 2-2).

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23

Figure 2-2. Front view and side view of the patella, quadriceps tendon (QT) and patellar tendon (PT) (Interactive Series, 1999).

Muscles

The primary muscles of the thigh can be classified into two groups: anterior and

posterior muscles according to their locations (Fig. 2-3). Of the anterior muscles, the

largest are the four muscles of the quadriceps femoris. The central rectus femoris

which is surrounded by the three vasti: The vastus intermedius, medialis, and lateralis.

Rectus femoris is attached to the pelvis with two tendons, while the vasti are inserted to

the femur. All four muscles unite in quadriceps tendon. The quadriceps is the knee

extensor. There are four posterior thigh muscles. The biceps femoris has two heads:

The long head has its origin on the ischial tuberosity together with the semitendinosus

and acts on two joints. The short head originates from the middle third of the linea

aspera on the shaft of the femur and the lateral intermuscular septum of thigh, and acts

on only one joint. These two heads unite to form the biceps which inserts on the head

of the fibula. The biceps flexes the knee joint and rotates the flexed leg laterally. It is

the only lateral rotator of the knee and thus has to oppose all medial rotator. The

semitendiosus and the semimembranosus share their origin with the long head of the

biceps, and both attaches on the medial side of the proximal head of the tibia together

with the gracilis and sartorius to form the pes anserinus. Functionally, the

semimembranosus and semitendinosus produce flexion and medial rotation at the

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knee.

Figure 2-3. Front view and back view of the main muscles around the knee joint (Interactive Series, 1999).

2.1.2 Functional Ankle and Foot Anatomy

The ankle and foot is a strong and complex mechanical structure containing 26 bones

(Fig. 2-4), 33 joints and over a hundred muscles, tendons and ligaments. These

structures together provide functions of support, balance and mobility to the body

through daily activities such as walking, running, and jumping. The foot can be

subdivided into the hindfoot, the midfoot, and the forefoot.

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Figure 2-4. Anatomy of the ankle and foot (Human kinetics publishers, Inc., 2012).

Hindfoot

The hindfoot is composed of the talus and the calcaneus. The tibia and fibula are

connected to the superior surface of the talus to form the talocrural joint. The calcaneus,

the largest bone of the foot, together with the inferior surface of talus, form the subtalar

joint. The bottom of the calcaneus is cushioned by the heel fat pad.

Midfoot

There are five irregular bones in the midfoot, including the cuboid, navicular and three

cuneiform bones. The joint between the hindfoot and the midfoot is called the midtarsal

joint. The midfoot is connected to the hindfoot and forefoot by muscles, ligaments and

the plantar fascia. The plantar fascia spans the entire length of the plantar surface,

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attaching the calcaneal tuberosity proximally and passing distally along the plantar

aspect of the foot, and dividing into five distal attachments at the base of the proximal

phalanges. It is composed mainly of longitudinally arranged fibers with a smaller

proportion of transversely arranged fibers. The tissue in central region of the plantar

fascia is thick and with a tensile strength larger the strength of the ligaments in the foot.

The plantar fascia plays an extremely important role in providing the stability and

support of the foot arches. When the plantar fascia is placed under excessive stress,

tears can occur, causing plantar fasciitis. Tension at one end of fascia is immediately

transmitted to the other, acting as the windlass mechanism. Lifting the toes tightens the

plantar fascia and deepens the foot arch (Fig. 2-5). The mechanism makes the foot

more rigid and more efficiently propel the body forward during walking. There are three

arches in the foot, including the medial longitudinal arch, lateral longitudinal arch and

transverse arch. The medial longitudinal arch (Fig. 2-5), the highest and most important

arch, is composed of the calcaneus, talus, navicular, medial cuneiform, and first

metatarsal bone. A flatfoot refers to a diminished medial longitudinal arch.

Figure 2-5. The medial longitudinal arch and windlass mechanism (Bandhayoga.com, Inc., 2012). The yellow band shows plantar fascia, the blue arc reveals foot arch.

Forefoot

The forefoot is composed of five toes and the corresponding five metatarsal bones.

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The bones of the toes are called phalanges. The joints between the phalange

components are called interphalangeal joint and those between the metatarsus and

phalanges are called metatarsophalangeal joint. Any of the three joints between the

tarsal and metatarsal bones are called tarsometatarsal joint, which link the midfoot to

the forefoot, including a medial joint between the first cuneiform and first metatarsal, an

intermediate joint between the second and third cuneiforms and corresponding

metatarsals, and a lateral joint between the cuboid and fourth and fifth metatarsals.

Muscles

All muscles in the leg are attached to the foot and can be classified into an anterior and

a posterior group. These two groups can be divided into subgroups or layers. The

anterior group consists of the extensors and the peroneus, while the posterior group

consists of a superficial and a deep layer. Functionally, the muscles of the leg are

either extensors or flexors. The extensors are responsible for the dorsiflexion of the

foot, while the flexors are responsible for the plantar flexion.

Three of the anterior muscles are extensors, as shown in Fig. 2-6. From its origin on

the lateral surface of the tibia, the tibialis anterior extends down to its insertion on the

plantar side of the medial cuneiform bone and the first metatarsal bone. In the non-

weight-bearing leg, the anterior tibialis dorsiflexes the foot and lifts the medial edge of

the foot. In the weight-bearing leg, it pulls the leg towards the foot. The tibialis anterior

is the main dorsiflexor of the ankle joint and invertor of the foot. It is most active during

the initial period of stance phase for slowing down ankle plantar flexion and preventing

the foot from slapping to the ground. It also prevents the forefoot from scraping the

ground during the swing phase.

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Figure 2-6. Side view and front view of the main muscles of the right leg (Interactive Series, 1999).

The extensor digitorum longus has a wide origin stretching from the lateral condyle of

the tibia down along the anterior side of the fibula. At the ankle, the tendon divides into

four that stretch across the foot to the dorsal aponeuroses of the last phalanges of the

four lateral toes. The extensor digitorum longus extends the lateral four toes at the

interphalangeal and metatarsophalangeal joints and assists in dorsiflexion of the ankle.

The extensor hallucis longus has its origin on the fibula and is inserted on the last

phalanx of hallux. The muscle dorsiflexes the hallux at the interphalangeal and

metatarsophalangeal joints and assists in dorsiflexion of the ankle.

Of the posterior muscles three are in the superficial layer, as shown in Fig. 2-7. The

major plantar flexors, commonly referred to as the triceps surae, are the gastrocnemius,

the medial head and lateral head of which arise on the distal end of the femur, and the

soleus, which arises on the proximal side of the tibia and fibula. The soleus muscle

acts with the gastrocnemius to plantarflex the ankle joint, preventing excessive

dorsiflexion during walking. These muscles together form the Achilles tendon, the

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largest and strongest tendon of the foot, which is attached to the posterior tubercle of

the calcaneus.

Figure 2-7. Back view of the main muscles of the right leg (from superficial to deep layer) (Interactive Series, 1999).

Tibialis posterior, flexor hallucis longus, flexor digitorum longus, peroneus longus and

peroneus brevis are the deep leg muscles, which arise from the tibia and fibula. All

these five muscles are weak ankle plantar flexors. The tibialis posterior originates from

the posterior surfaces of the shafts of the tibia and fibula and runs down behind the

medial malleolus. It inserts onto the tuberosity of the navicular and extends to the

second, third and fourth metatarsal bases. The muscle produces simultaneous plantar

flexion and supination of the foot.

The flexor hallucis longus (Fig. 2-7) arises distally on the fibula. Its tendon runs forward

and attaches on the plantar surface of the distal phalanx of the hallux. It plantarflexes

the hallux and assists in supination. The flexor hallucis longus acts during final

propulsion in gait and assists in the maintenance of the medial longitudinal arch. The

flexor digitorum longus (Fig. 2-7) has its origin on the upper part of the tibia and divides

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into four tendons attached to the last phalanges of the four lateral toes. In the non-

weight-bearing leg, it plantar flexes the toes and supinates the foot. In the weight-

bearing leg, it supports the plantar arch.

Two muscles on the lateral side of the leg form the peroneus group. The peroneus

longus and brevis have their origins on the fibula and they both pass behind the lateral

malleolus. Beneath the foot, the peroneus longus stretches from the lateral to the

medial side in a groove, thus bracing the transverse arch of the foot. The peroneus

brevis is attached on the lateral side to the tuberosity of the fifth metatarsal. Together

the two peroneus form the strongest pronators of the foot and stabilize the subtalar and

midtarsal joints for propulsion.

2.1.3 Biomechanical factors to knee joint degenerations

Due to the forces and moments acting on the knee joint in terms of both acute

traumatic impacts and chronic weight bearing problems, the knee sustains high joint

loadings and is vulnerable to lesions stimulated by mechanical factors. Knee joint

together with the ankle joint contribute a lot to the lower extremity structure such as

alignment and stability. Although the causative mechanisms of knee problems are not

entirely legible, increased joint loadings during walking have been associated with the

initiation and progression of knee pain (Andriacchi and Mundermann, 2006).

The knee alignment, as shown by the hip/knee/ankle angle, is an important

determinant of loading distributions in the knee joi

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