Foot and Ankle Biomechanics 9780128154496, 0128154497

Table of contents :
Cover
Foot and Ankle Biomechanics
Copyright
List of contributors
Dedication
Contents
Preface
Anatomical Terms Used in Foot and Ankle Biomechanics
1 Anatomy of the Foot
1.1 Skeletal structures
1.1.1 Tibia and fibula
1.1.2 Segments of the foot skeleton
1.1.3 Talus
1.1.4 Calcaneus
1.1.5 Navicular
1.1.6 Cuboid
1.1.7 Medial cuneiform
1.1.8 Intermediate cuneiform
1.1.9 Lateral cuneiform
1.1.10 Metatarsals
1.1.10.1 First metatarsal
1.1.10.2 Second metatarsal
1.1.10.3 Third metatarsal
1.1.10.4 Fourth metatarsal
1.1.10.5 Fifth metatarsal
1.1.11 Phalanges
1.1.11.1 Proximal phalanges
1.1.11.2 Middle phalanges
1.1.11.3 Distal phalanges
1.2 Joints
1.2.1 Tibiofibular syndesmosis
1.2.2 Ankle joint
1.2.3 Subtalar joint
1.2.4 Talocalcaneonavicular joint
1.2.5 Calcaneocuboid joint
1.2.6 Cuboideonavicular joint
1.2.7 Cuneonavicular joints
1.2.8 Intercuneiform and cuneocuboid joints
1.2.9 Tarsometatarsal joints
1.2.10 Proximal intermetatarsal joints
1.2.11 Distal intermetatarsal joints
1.2.12 Lesser metatarsophalangeal joints
1.2.13 Hallucal metatarsophalangeal joint
1.2.14 Interphalangeal joints
1.3 Muscles and fascial specializations
1.3.1 Fascial specializations
1.3.2 Extrinsic dorsal muscles
1.3.2.1 Tibialis anterior
1.3.2.2 Extensor digitorum longus
1.3.2.3 Extensor hallucis longus
1.3.2.4 Peroneus tertius
1.3.3 Extrinsic plantar muscles
1.3.3.1 Triceps surae
1.3.3.2 Plantaris
1.3.3.3 Flexor digitorum longus
1.3.3.4 Flexor hallucis longus
1.3.3.5 Tibialis posterior
1.3.4 Extrinsic lateral muscles
1.3.4.1 Peroneus longus
1.3.4.2 Peroneus brevis
1.3.5 Intrinsic dorsal foot muscles
1.3.5.1 Extensor hallucis brevis and extensor digitorum brevis
1.3.6 Intrinsic plantar muscles
1.3.6.1 Abductor hallucis
1.3.6.2 Flexor digitorum brevis
1.3.6.3 Abductor digiti minimi
1.3.6.4 Quadratus plantae
1.3.6.5 Lumbricals
1.3.6.6 Flexor hallucis brevis
1.3.6.7 Flexor digiti minimi brevis
1.3.6.8 Adductor hallucis
1.3.6.9 Dorsal and plantar interossei
1.4 Nerves
1.4.1 The fibular nerves in the foot
1.4.2 The tibial nerves in the foot
1.5 Blood supply
1.5.1 Arteries
1.5.2 Veins
1.5.2.1 Superficial veins
1.5.2.2 Deep veins
1.5.3 Lymphatics
Further reading
2 Basic Biomechanics
2.1 Introduction
2.2 Terminology
2.3 Statics
2.4 Dynamics
2.5 Strength of materials and deformation
2.6 Viscoelasticity
2.7 Summary
References
3 Anatomical Nomenclature: Conundrums of Nonstandardized Foot and Ankle Terminology
3.1 Introduction
3.2 Anatomical descriptions
3.2.1 The persistence of eponyms
3.2.2 Regional descriptions
3.2.2.1 The distal element of the lower limb
3.2.2.2 Forefoot and hindfoot
3.2.3 Applying the anatomical position
3.2.4 Application specific human anatomical positions
3.2.5 Defining anatomical directions, planes, and axes
3.2.5.1 Posterior–anterior versus ventral–dorsal
3.2.5.2 Anatomical planes
3.2.5.3 Foot midline
3.3 Foot motions
3.3.1 Defining motions
3.3.1.1 Flexion-extension
3.3.1.2 Adduction-abduction versus external-internal rotation
3.3.2 Whole foot motions and their complexity
3.3.2.1 Plantarflexion-dorsiflexion
3.3.2.2 Inversion—eversion
3.3.2.3 Pronation—supination
3.4 Terminological implications of mathematical choices
3.5 Conclusion: standardizing foot and ankle terminology
References
4 Kinematics and Kinetics of the Foot and Ankle during Gait
4.1 Introduction
4.2 Overview of relevant anatomy
4.3 Overview of kinematic and kinetic modeling
4.4 Healthy and impaired feet
4.5 Multisegment foot models
4.6 Future areas of research
4.6.1 Biplane fluoroscopy
4.6.2 Modeling
4.7 Conclusion
References
5 Bone, Cartilage, and Joint Function
5.1 Bone components and structure
5.2 Cartilage
5.3 Joint functions
5.3.1 Talocrural joint
5.3.2 Talocalcaneal (subtalar) joint
5.3.3 Transverse tarsal joint
5.3.4 Tarsometatarsal joint
5.3.5 Metatarsophalangeal joint
5.4 Areas of future research
References
6 Muscles and Tendons
6.1 Introduction
6.2 Biomechanical function
6.2.1 Normal foot
6.2.1.1 Foot stability
6.2.1.2 Balance
6.2.1.3 Locomotion
6.2.2 Aging
6.2.3 Pathologies
6.2.3.1 Plantar fasciitis
6.2.3.2 Pes planus
6.2.3.3 Toe deformities
6.2.3.4 Diabetic neuropathy
6.2.4 Footwear and orthoses
6.2.5 Interventions
6.3 Areas of future biomechanical research
References
7 Ligaments
7.1 Introduction
7.2 Ligament anatomy
7.3 Mechanical properties
7.3.1 Ankle joint
7.3.1.1 Lateral collateral ligaments
7.3.1.2 Medial collateral ligaments
7.3.2 Hindfoot ligaments
7.3.2.1 Subtalar joint
7.3.2.2 Talonavicular joint
7.3.3 Midfoot ligaments
7.3.3.1 Plantar fascia
7.3.3.2 Metatarsal base ligaments
7.3.4 Forefoot
7.3.5 Variations in mechanical properties
7.3.5.1 Changes in activity level
7.3.5.2 Foot comorbidity
7.3.5.3 Age effects
7.3.5.4 Influence of anthropometric effects
7.4 Ligament sprains
7.5 Overcoming limitations
7.6 Future areas of research
References
8 Plantar Soft Tissue
8.1 Introduction
8.2 Anatomy
8.2.1 Gross anatomy
8.2.2 Histological or biochemical
8.2.3 Medical imaging of tissue thickness
8.3 Biomechanical function
8.4 Mechanical properties
8.4.1 Structural in vivo testing
8.4.2 Structural ex vivo testing
8.4.3 Material ex vivo testing
8.4.4 Ultrasound
8.4.5 Other in vivo techniques
8.5 Effect of aging
8.6 Diabetic plantar soft tissue
8.6.1 Other pathologies associated with the plantar soft tissue
8.7 Areas of future biomechanical research
References
9 Multisegment Foot Models
9.1 Basic principles of multisegment foot models
9.1.1 Overview of motion capture
9.1.2 Why use a multisegment foot model?
9.2 Selecting an appropriate multisegment foot model
9.2.1 Segments and bone groupings
9.2.2 Marker type and placement
9.2.3 Coordinate systems
9.2.4 Offsets
9.2.5 Standardized description of multisegment foot models
9.3 Review of current multisegment foot models
9.3.1 Milwaukee kinematic model
9.3.2 Leardini/Rizzoli kinematic model
9.3.3 Oxford kinematic model
9.3.4 MacWilliams/Kinfoot kinetic model
9.3.5 Bruening kinetic model
9.3.6 Direct comparison of current multisegment foot models
9.4 Applications, considerations, and limitations
9.4.1 Clinical applications
9.4.2 Sources of error
9.5 Areas of future biomechanical research
References
10 Invasive Techniques for Studying Foot and Ankle Kinematics*
10.1 Introduction
10.2 Early invasive studies of foot and ankle biomechanics
10.3 Radiostereometric analysis
10.3.1 Technique
10.3.2 Talocrural joint
10.3.3 Relationship between the joints distal to the talus
10.3.4 Transferral of rotation between the leg and the foot
10.3.5 Ankle mortise width
10.4 Applications and significance of studies using intracortical pins for foot and ankle kinematics
10.4.1 Skin movement artifact in foot and ankle kinematics
10.4.2 Ankle kinematics
10.4.3 Foot and ankle basic research in walking and running kinematics
10.4.4 Applied studies of orthoses and shoe conditions
10.5 Limitations and future directions
Appendix: Insertion of markers in bones of the foot and ankle
References
11 Biplane Fluoroscopy
11.1 Introduction
11.2 Background and history of biplane fluoroscopy
11.2.1 Overview of how X-ray imaging works
11.2.2 Overview of biplane system history and evolution
11.2.2.1 Intact C-arm systems for foot bone tracking
11.2.2.2 Disarticulated C-arm systems for foot bone tracking
11.2.2.3 Custom dedicated biplane hardware for foot bone tracking
11.3 Other techniques for tracking foot bone kinematics
11.4 Challenges specific to foot and ankle tracking with biplane fluoroscopy
11.5 Overview of biplane hardware
11.6 Overview of biplane software
11.7 Clinical biplane foot and ankle studies
11.7.1 Biplane systems consisting of two C-arms
11.7.2 Biplane systems consisting of two modified C-arms
11.7.3 Biplane systems consisting of independent X-ray sources and image intensifiers
11.8 Future applications and directions
References
12 Plantar Pressure and Ground Reaction Forces
12.1 Introduction: clinical relevance of force and pressure measurements in foot and ankle biomechanics
12.2 Background: force versus pressure
12.2.1 History
12.2.2 Development of measurement technologies
12.2.3 Visualization and analytical options
12.3 Research applications and selected clinical examples
12.3.1 Diabetes
12.3.2 Children’s flatfoot
12.3.3 Sports
12.4 Areas of future research
References
13 Electromyography and Dynamometry for Investigating the Neuromuscular Control of the Foot and Ankle
13.1 Introduction
13.2 Electromyography
13.2.1 Surface electromyography
13.2.2 Indwelling electromyography and motor unit recordings
13.3 Dynamometry
13.3.1 Isometric
13.3.2 Dynamic
13.4 Ankle and foot related considerations and insights
13.4.1 Motor unit behavior and quantity
13.4.2 Maximal voluntary contractions and knee angle
13.4.3 History-dependence of force
13.5 Future research
References
14 From Impossible to Unnoticed: Wearable Technologies and The Miniaturization of Grand Science
14.1 Introduction
14.1.1 Tech affords understanding
14.1.2 Wearable domains
14.1.3 Breakout box: what makes a successful wearable device?
14.2 The past
14.2.1 Force and pressure
14.2.2 Health and activity sensing
14.3 The present
14.3.1 Force and pressure sensing
14.3.2 Health and activity monitoring
14.3.3 Actuation and assistance
14.3.4 Haptics
14.3.4.1 Sensory substitution
14.3.4.2 Cueing and notification
14.3.5 Motion capture
14.4 The future
References
15 Integrated Laboratories for Pursuing Pedal Pathologies
15.1 Introduction
15.2 Our method of approach
15.3 Integrated laboratories
15.3.1 Epidemiology
15.3.2 In vivo experimentation
15.3.2.1 Gait analysis
15.3.2.2 Plantar pressure assessments
15.3.2.3 Measures of foot structure
15.3.2.4 Other measures
15.3.3 In vitro experimentation
15.3.3.1 Histology
15.3.3.2 Cadaveric simulators
15.3.4 In silico simulation
15.3.4.1 Medical image processing
15.3.4.2 Finite element modeling
15.3.4.3 Musculoskeletal modeling
15.3.4.4 Sensitivity studies
15.3.4.5 Validation
15.4 Case study of the integrated laboratories concept to the study of hallux rigidus
15.4.1 Epidemiology
15.4.2 In vivo experimentation
15.4.3 In vitro experimentation
15.4.4 In silico simulation
15.5 Future biomechanical research
References
16 Radiographs
16.1 Introduction
16.2 Radiographic technology
16.3 Standard radiographic views of the foot and ankle
16.4 Definitions of X-ray measurements of foot shape
16.5 Foot-specific applications and considerations
16.6 Clinical X-ray measures of foot shape
16.7 Issues with X-ray measures of foot shape
16.8 Areas of future biomechanical research
References
17 Computed Tomography of the Foot and Ankle
17.1 Introduction
17.1.1 History and development of computed tomography
17.1.2 Comparison to other imaging modalities
17.1.3 Computed tomography protocols for the foot and ankle
17.2 Foot-specific applications and considerations
17.2.1 Disease diagnosis
17.2.2 Surgical assessment and planning
17.2.3 Biomechanics research
17.2.3.1 Kinematic measurements
17.2.3.2 Bone density properties
17.2.3.3 Computational models
17.2.3.4 Shape modeling and assessment
17.3 Areas of future biomechanical research
References
18 Weight-bearing Computed Tomography of the Foot and Ankle
18.1 Introduction
18.2 Biases of conventional radiography
18.3 Technical aspects
18.4 Indications
18.5 3D biometrics
18.6 Advantages and limitations of weight-bearing computed tomography
18.7 Future areas of research
18.8 Conclusion
Acknowledgments
Conflict of interest statement
References
Further reading
19 Magnetic Resonance Imaging of the Foot and Ankle
19.1 Introduction
19.2 Magnetic resonance imaging sequences
19.3 Magnetic resonance imaging versus computed tomography
19.4 Magnetic resonance imaging appearance of musculoskeletal tissue—normal and pathology
19.4.1 Short T2 tissues
19.4.2 Tendons
19.4.3 Pseudo tendon pathology—the magic angle effect or phenomenon
19.4.4 Ligaments
19.4.5 Bone
19.4.6 Hyaline and fibrocartilaginous cartilage
19.4.7 Muscle
19.4.8 Bursa/synovia
19.5 Tailored magnetic resonance imaging protocol for the foot and ankle—indication driven
19.5.1 Imaging orientation
19.5.2 Tailored magnetic resonance imaging protocols
19.5.3 Optimized imaging planes
19.5.4 Metal artifact reduction sequences
19.6 Magnetic resonance imaging anatomy of the foot and ankle
19.6.1 Ankle
19.6.1.1 Ligaments
19.6.1.2 Bony defects and ligament injuries
19.6.2 Hindfoot
19.6.2.1 Tendons
19.6.2.2 Enthesitis—plantar fasciitis
19.6.3 Midfoot
19.6.3.1 Degenerative joint disease
19.6.3.2 Stress fractures
19.6.4 Forefoot
19.6.4.1 Morton’s neuroma
19.6.4.2 Osteomyelitis
19.6.4.3 Plantar plate tears
19.7 Areas of future research
19.7.1 Radiation-free bone imaging
19.7.2 Magnetic resonance imaging scan time reduction
19.7.3 Volumetric, isotropic 3D magnetic resonance imaging sequences
References
20 Biomechanical Assessment of Soft Tissues in the Foot and Ankle Using Ultrasound
20.1 Introduction
20.1.1 Ultrasound imaging, how it works
20.2 Ultrasound assessment of structural changes: the effect of weightbearing activities
20.2.1 Assessment of plantar soft tissue thickness in relation to weightbearing activities
20.2.2 Assessment of plantar fascia thickness in relation to weightbearing activities
20.2.3 Assessment of Achilles tendon thickness in relation to weightbearing activities
20.2.4 Summary and limitations of weightbearing ultrasound
20.3 Ultrasound assessment combined with measurement of load
20.3.1 Ultrasound combined with load cells to assess the mechanical properties of the plantar soft tissue
20.3.2 Ultrasound combined with dynamometry to assess the mechanical properties of Achilles tendon
20.3.3 Summary and limitations of ultrasound assessment combined with measurement of load
20.4 Ultrasound elastography (sonoelastography)
20.4.1 Ultrasound elastography to assess the mechanical properties of plantar soft tissue
20.4.2 Ultrasound elastography to assess the mechanical properties of plantar fascia
20.4.3 Ultrasound elastography to assess the mechanical properties of Achilles tendon
20.4.4 Summary and limitations of ultrasound elastography
20.5 Conclusion and future areas of research
References
21 3D Surface Scanning of the Foot and Ankle
21.1 Introduction
21.1.1 The history and development of 3D scanning
21.1.2 Technologies
21.1.2.1 Contact scanners (coordinate measuring machines)
21.1.2.2 Noncontact scanners
21.2 Foot-specific applications and considerations
21.2.1 Reliability and comparisons to other techniques
21.2.2 Population studies
21.2.3 Orthoses and footwear
21.2.4 Musculoskeletal models
21.2.5 Bones and other internal anatomy
21.3 Areas of future biomechanical research
References
22 Cadaveric Gait Simulation
22.1 Introduction
22.2 Techniques for dynamic gait simulation
22.3 Limitations of dynamic gait simulation
22.4 Clinical applications of dynamic gait simulation
22.5 Areas of future biomechanical research
22.6 Conclusion
References
23 Finite Element Modeling
23.1 Introduction
23.2 Basic concepts of finite element modeling
23.3 Applications of finite element analysis in foot biomechanics
23.3.1 Simulation of the interaction between foot and footwear
23.3.2 Simulation of healthy foot biomechanics
23.3.3 Finite element modeling for the in vivo material characterization of soft tissues
23.3.4 Simulation of pathological conditions
23.3.4.1 Biomechanics of the pathologic foot
23.3.4.2 Study of surgical interventions
23.4 Modeling strategies
23.4.1 Geometry design
23.4.1.1 3D modeling versus 2D modeling
23.4.1.2 Modeling of the entire foot compared to anatomically focused modeling
23.4.1.3 Anatomically detailed compared to idealized modeling
23.4.2 Meshing
23.4.2.1 Element type selection
23.4.2.2 Mesh convergence
23.4.3 Material properties
23.4.3.1 Bone and cartilage
23.4.3.2 Ligaments and tendons
23.4.3.3 Soft tissues
23.4.4 Solver selection
23.4.5 Reliability assessment
23.5 Limitations and future research toward clinically applicable finite element modeling
23.6 Summary
References
24 Musculoskeletal Modeling of the Foot and Ankle
24.1 Introduction
24.1.1 Musculoskeletal models and their development
24.1.2 Validation techniques
24.1.3 Challenges in modeling the foot and ankle
24.2 Foot specific models and applications
24.3 Areas of future biomechanical research
References
25 Predicting and Preventing Posttraumatic Osteoarthritis of the Ankle
25.1 Introduction: pathomechanical origins of posttraumatic osteoarthritis
25.2 Pathomechanics I: acute joint injury severity
25.3 Pathomechanics II: chronic stress aberration
25.4 Pathomechanics III: altered kinematics
25.5 Areas of future biomechanical research
25.5.1 Posttraumatic ankle osteoarthritis: opportunities for intervention informed by pathomechanical knowledge
25.6 Summary/conclusion
References
26 Mechanics of Biological Tissues
26.1 Introduction
26.2 Materials and methods
26.2.1 Finite element modeling of the foot
26.2.2 Formulation of constitutive models
26.2.2.1 Linear elastic constitutive models
26.2.2.2 Hyperelastic constitutive models
26.2.2.3 Visco-hyperelastic model
26.2.3 Identification of constitutive parameters
26.2.3.1 Constitutive parameter identification for bone
26.2.3.2 Constitutive parameter identification for cartilage
26.2.3.3 Constitutive parameter identification for plantar soft tissue
26.2.3.4 Constitutive parameter identification for ankle ligaments
26.2.4 Numerical analyses of foot functionality
26.2.4.1 Ankle movements
26.2.4.2 Gait cycle
26.2.4.3 Foot and footwear interaction
26.2.4.4 Diabetic condition
26.2.5 Limitations of computational modeling
26.2.6 Future biomechanics research
26.3 Conclusion
References
27 Clinical Examination of the Foot and Ankle
27.1 Introduction
27.2 Demographics
27.3 Vital signs
27.4 Patient history
27.5 Assessment of pain
27.6 Visual observation/inspection
27.6.1 Skin
27.6.2 Edema
27.6.3 Atrophy
27.6.4 Temperature
27.6.5 Scarring
27.6.6 Callus patterns
27.6.7 Exostoses
27.6.8 Ankle and foot deformities
27.7 Lower extremity alignment
27.8 Foot posture or foot shape
27.8.1 Planus foot type
27.8.2 Cavus foot type
27.9 Limb length
27.10 Radiographic examination
27.11 Range of motion/flexibility/joint mobility
27.12 Joint mobility
27.13 Ligamentous/stability testing
27.14 Tendon
27.15 Muscle strength
27.16 Sensory testing
27.17 Circulation
27.18 Foot and ankle specific testing
27.19 Footwear examination
27.20 Functional assessment
27.21 Outcomes assessment
27.22 Areas of future biomechanical research
References
28 Foot Type Biomechanics
28.1 Introduction
28.2 Structural foot type
28.3 Functional foot type
28.4 Foot type biomechanics
28.5 Association with pain and injury
28.5.1 Pain and injury
28.6 Treatments
28.7 Areas of future biomechanical research
References
29 Traumatic Foot and Ankle Injuries
29.1 Introduction
29.2 Pilon fractures
29.2.1 Etiology and pathophysiology
29.2.2 Symptoms
29.2.3 Diagnostics/classification
29.2.4 Treatment
29.3 Calcaneal fractures
29.3.1 Etiology and pathophysiology
29.3.2 Symptoms
29.3.3 Diagnostics/classification
29.3.4 Treatment
29.4 Talus fractures
29.4.1 Etiology and pathophysiology
29.4.2 Symptoms
29.4.3 Diagnostics/classification
29.4.4 Treatment
29.5 Tarsometatarsal (Lisfranc) injuries
29.5.1 Etiology and pathophysiology
29.5.2 Symptoms
29.5.3 Diagnostics/classification
29.5.4 Treatment
29.6 Metatarsal fractures
29.6.1 Etiology and pathophysiology
29.6.2 Symptoms
29.6.3 Diagnostics/classification
29.6.4 Treatment
29.7 Midfoot crush injuries
29.7.1 Etiology and pathophysiology
29.7.2 Symptoms
29.7.3 Treatment
29.8 Acute ankle sprains
29.8.1 Etiology and pathophysiology
29.8.2 Symptoms
29.8.3 Treatment
29.9 Syndesmosis tears
29.9.1 Etiology and pathophysiology
29.9.2 Symptoms
29.9.3 Treatment
29.10 Achilles tendon rupture
29.10.1 Etiology and pathophysiology
29.10.2 Symptoms
29.10.3 Treatment
29.11 Areas of future research
References
30 The Pediatric Foot
30.1 Introduction
30.2 Common pathologies affecting pediatric feet
30.2.1 Congenital foot deformities
30.2.2 Developmental foot deformities
30.2.3 Foot pathologies associated with other conditions
30.3 Functional assessment of the pediatric foot
30.4 Areas for future research
References
31 Neurological Foot Pathology
31.1 Introduction
31.2 Stroke
31.2.1 Pathology related to the musculoskeletal system
31.2.2 Impact on kinematics
31.2.3 Impact on foot function
31.2.4 Clinical treatment
31.3 Cerebral palsy
31.3.1 Definition
31.3.2 Structural deformities and gait deviations
31.3.3 Treatment
31.4 Toe walking
31.4.1 Diagnosis and etiology
31.4.2 Biomechanical and musculoskeletal function
31.4.3 Treatment
31.5 Peripheral neuropathy
31.5.1 Background
31.5.2 Musculoskeletal and movement implications
31.6 Foot drop
31.6.1 Pathology
31.6.2 Impact on biomechanics
31.6.3 Treatment
31.7 Tarsal tunnel syndrome
31.7.1 Pathology
31.7.2 Impact on biomechanics
31.7.3 Treatment
31.8 Morton’s neuroma
31.8.1 Pathology
31.8.2 Impact on biomechanics
31.8.3 Treatment
31.9 Charcot foot
31.9.1 Pathology
31.9.2 Impact on biomechanics
31.9.3 Treatment
31.10 Charcot-Marie-Tooth disease
31.10.1 Pathology
31.10.2 Impact on biomechanics—pediatric and young adult
31.10.3 Impact on biomechanics
31.10.4 Treatment
31.11 Friedreich’s ataxia
31.11.1 Background and pathology
31.11.2 Gait analysis
31.11.3 Musculoskeletal effects
31.11.4 Clinical treatment
31.12 Poliomyelitis
31.12.1 Pathology
31.12.2 Current status
31.12.3 Impact on biomechanics
31.12.4 Treatment
31.13 Areas of Future Research
References
32 Chronic Foot and Ankle Injuries
32.1 Introduction
32.1.1 Chronic injury through microtrauma
32.1.2 Chronic injury through macrotrauma
32.1.3 Impairment-based rehabilitation model for treating chronic injuries
32.1.4 Role of patient-oriented outcomes
32.2 Chronic ankle instability
32.2.1 Anatomy overview
32.2.2 Etiology
32.2.2.1 Mechanism of injury and pathomechanics
32.2.3 Clinical impairments
32.2.3.1 Range of motion
32.2.3.2 Strength
32.2.3.3 Balance
32.2.3.4 Functional activity
32.2.4 Treatment
32.2.4.1 Acute management
32.2.4.2 On-going management
32.3 Plantar fasciitis
32.3.1 Anatomical overview
32.3.2 Etiology
32.3.2.1 Mechanism of injury and pathomechanics
32.3.3 Clinical impairments
32.3.3.1 Range of motion
32.3.3.2 Strength
32.3.3.3 Functional activity
32.3.4 Treatment
32.3.4.1 Acute management
32.3.4.2 On-going management of clinical impairments
32.4 Tendinopathy (Achilles, peroneal, and posterior tibialis)
32.4.1 Anatomical overview
32.4.2 Etiology
32.4.2.1 Mechanism of injury and pathomechanics
32.4.2.1.1 Achilles tendon
32.4.2.1.2 Tibialis posterior tendon
32.4.2.1.3 Peroneal tendon
32.4.3 Clinical impairments
32.4.4 Treatment
32.5 Stress fractures (navicular, metatarsals)
32.5.1 Anatomical overview
32.5.2 Etiology
32.5.2.1 Mechanism of injury and pathomechanics
32.5.3 Clinical impairments
32.5.4 Treatment
32.6 Sesamoiditis
32.6.1 Anatomical overview
32.6.2 Etiology
32.6.2.1 Mechanism of injury and pathomechanics
32.6.3 Clinical impairments
32.6.4 Treatment
32.7 Retrocalcaneal bursitis
32.7.1 Anatomical overview
32.7.2 Etiology
32.7.2.1 Mechanism of injury and pathomechanics
32.7.3 Clinical impairments
32.7.4 Treatment
32.8 Areas of future research for chronic foot and ankle injuries
References
33 Hallux Valgus
33.1 Introduction
33.2 Prevalence
33.3 Etiology
33.3.1 Genetics and race
33.3.2 Structural and biomechanical factors
33.3.3 Footwear
33.4 Diagnosis and imaging
33.4.1 Clinical diagnosis
33.4.2 Radiographic assessment
33.4.3 Ultrasound
33.4.4 Computed tomography
33.4.5 Magnetic resonance imaging
33.5 Clinical presentation
33.5.1 Foot pain
33.5.2 Footwear
33.5.3 Self-reported function and quality of life
33.6 Functional outcomes
33.6.1 Balance and falls
33.6.2 Hallux flexion and abduction
33.6.3 Gait analysis
33.6.3.1 Kinematics
33.6.3.2 Plantar pressures
33.6.3.3 Muscle activity
33.6.3.4 Temporospatial parameters
33.7 Treatment pathways
33.7.1 Nonsurgical treatment
33.7.1.1 Expert opinion and current practice
33.7.1.2 Foot orthoses
33.7.1.3 Splints and toe separators
33.7.1.4 Manual therapy
33.7.1.5 Taping
33.7.1.6 Exercise
33.7.1.7 Botulinum toxin A injection
33.7.2 Surgical treatment
33.8 Future directions for research
33.9 Summary
References
34 Osteoarthritis of the Foot and Ankle
34.1 Introduction
34.1.1 Osteoarthritis symptoms and diagnosis
34.1.2 Structural changes
34.1.3 Risk factors and classification of osteoarthritis
34.1.4 Foot and ankle osteoarthritis subtypes
34.2 First metatarsophalangeal joint osteoarthritis
34.2.1 Etiology and impact
34.2.2 Clinical findings
34.2.3 Structural and biomechanical features
34.2.4 Clinical and biomechanical effects of conservative treatment
34.3 Midfoot osteoarthritis
34.3.1 Etiology and impact
34.3.2 Clinical findings
34.3.3 Structural and biomechanical features
34.3.4 Clinical and biomechanical effects of conservative treatment
34.4 Ankle osteoarthritis
34.4.1 Etiology and impact
34.4.2 Clinical findings
34.4.3 Structural and biomechanical features
34.4.4 Clinical and biomechanical effects of conservative treatment
34.5 Areas of future biomechanical research
34.6 Summary
References
35 Diabetic Foot Disease
35.1 Background on diabetes
35.2 Overview of key negative outcomes of diabetic foot disease
35.2.1 Diabetic peripheral neuropathy
35.2.2 Peripheral vascular disease
35.2.3 Diabetic plantar ulceration
35.2.4 Foot deformities
35.2.5 Lower-extremity fractures
35.2.6 Charcot neuroarthropathy
35.2.7 Lower extremity amputation
35.3 Risk factors for the development and progression of diabetic foot disease
35.4 Changes in kinematics and kinetics in diabetic foot disease
35.5 Changes in tissue characteristics
35.5.1 Muscle: fatty infiltration and reduction of intrinsic foot muscle volumes
35.5.2 Bone
35.5.3 Cartilage
35.5.4 Tendon
35.5.5 Plantar fascia
35.6 The relationship between foot deformities and plantar ulceration
35.7 The relationship between lower extremity fractures and Charcot neuropathic osteoarthropathy
35.8 Areas of future biomechanical research
References
36 Rheumatic Foot Disease
36.1 Introduction
36.2 Rheumatoid arthritis
36.2.1 Early rheumatoid arthritis
36.2.2 Established rheumatoid arthritis
36.3 Spondlyarthropathies
36.4 Juvenile idiopathic arthritis
36.5 Connective tissue disorders
36.6 Gout
36.7 Future research
References
37 The Aging Foot
37.1 Changing properties and functions of foot tissues
37.1.1 Bone
37.1.2 Cartilage
37.1.3 Muscle
37.1.4 Tendon
37.1.5 Ligament
37.1.6 Skin
37.1.7 Neural
37.1.8 Fat pad
37.2 Foot posture and morphology
37.2.1 Anthropometrics
37.2.2 Foot posture
37.2.3 Arch height
37.2.4 Joint range-of-motion
37.3 Foot function (kinematics/kinetics/plantar pressures)
37.3.1 Kinetics
37.3.1.1 Ground reaction forces
37.3.1.2 Joint moments and powers
37.3.1.3 Plantar pressures
37.3.2 Kinematics
37.3.2.1 Ankle and foot
37.4 Foot posture, foot disorders, and mobility limitations
37.4.1 Foot posture and foot deformity
37.4.2 Foot posture and foot symptoms
37.4.3 Foot posture, mobility limitations, and falls
37.5 Areas for future research
References
38 Biomechanics of Athletic Footwear
38.1 Introduction
38.2 Anatomy of a running shoe
38.3 Biomechanics of athletic footwear design
38.3.1 Cushioning
38.3.2 Hindfoot stability
38.4 Types of shoes and their features
38.4.1 Casual shoes
38.4.2 Running shoes
38.4.3 Racing flats & spikes
38.4.4 Marathon shoes
38.4.5 Other sports shoes
38.4.6 New shoe innovations
38.4.6.1 Footwear embedded energy harvester
38.4.6.2 Lacing systems
38.4.6.3 3D printed midsoles and outsoles
38.4.7 Graphene outsoles
38.5 Shod versus barefoot
38.6 Footwear related injuries
38.7 Future footwear research
References
39 Minimal Shoes: Restoring Natural Running Mechanics
39.1 Introduction
39.2 Brief history of running footwear
39.3 Biomechanics of barefoot and conventional shod running
39.3.1 Barefoot running pattern
39.3.2 Conventional shod running pattern
39.3.3 Comparison of mechanics between conventional shod and barefoot running
39.4 Minimal footwear running
39.4.1 Definition of full minimalist footwear
39.4.2 Comparison of full minimal to barefoot running
39.4.3 Comparison of full minimal to partial minimal shoes
39.4.4 Comparison of full minimal to conventional footwear
39.5 Effect of minimal shoes on the foot musculoskeletal system
39.6 Summary
39.7 Future research
References
40 Foot Orthoses
40.1 Introduction
40.1.1 Design and manufacture of foot orthoses
40.2 Biomechanical effects of foot orthoses
40.2.1 Kinematic effects of foot orthosis
40.2.2 Kinetic effects of foot orthosis
40.2.3 Effects of foot orthosis on plantar pressure
40.2.4 Effects of foot orthosis on muscle activity patterns
40.3 Effects of foot orthosis on clinical conditions
40.3.1 Rheumatoid arthritis
40.3.2 Symptomatic flat foot
40.3.3 Heel pain (plantar fasciitis)
40.3.4 Osteoarthritis
40.3.5 Sports injuries and other conditions
40.4 Areas of future research
References
41 Ankle-Foot Orthoses and Rocker Bottom Shoes
41.1 Introduction
41.2 Ankle-foot orthoses
41.2.1 Controlling rotational motion
41.2.2 Controlling translational motion
41.2.3 Controlling axial forces
41.2.4 Altering the line of action of the ground reaction force
41.3 Rocker bottom shoes
41.4 Roll-over shape
41.5 Patient populations
41.5.1 Stroke
41.5.2 Cerebral palsy
41.5.3 Ankle arthritis
41.5.4 Limb salvage
41.6 Design and prescription of ankle-foot orthosis
41.6.1 Conventional versus advanced
41.6.2 Articulated versus nonarticulated
41.7 Design and prescription of rocker bottom shoes
41.8 Variations on materials
41.9 New designs
41.10 Sport applications
41.11 Areas of future research
References
42 Diabetic Footwear
42.1 Introduction
42.2 Foot biomechanics and offloading
42.3 The biomechanical effect of diabetic footwear and offloading devices
42.4 Footwear and offloading for ulcer healing
42.5 Diabetic footwear for ulcer prevention
42.6 Footwear and offloading adherence
42.7 Other considerations
42.8 Future research
42.9 Conclusions
References
43 Reconstructions for Adult-acquired Flatfoot Deformity
43.1 Introduction
43.2 Hindfoot valgus
43.2.1 Bony anatomy
43.2.2 Ligament failure
43.2.3 Surgical reconstruction
43.3 Forefoot external rotation
43.3.1 Bony anatomy
43.3.2 Ligament and tendon failure
43.3.3 Surgical reconstruction
43.4 Sag at the talonavicular joint
43.4.1 Bony anatomy
43.4.2 Ligamentous failure
43.4.3 Surgical reconstruction
43.5 Failure of the posterior tibial tendon
43.5.1 Clinical assessment
43.5.2 MRI assessment
43.5.3 Surgical reconstruction
43.6 Gastrocnemius and Achilles tightness
43.6.1 Clinical assessment
43.6.2 Surgical treatment
43.7 Medial arch eversion
43.7.1 Clinical assessment
43.7.2 Bony anatomy
43.7.3 Ligamentous failure
43.7.4 Surgical reconstruction
43.8 Special considerations
43.9 Future biomechanical studies and conclusion
References
44 Cavus Foot Reconstructions
44.1 Introduction
44.2 Etiology
44.2.1 Traumatic
44.2.2 Neurologic
44.2.3 Residual clubfoot
44.2.4 Idiopathic
44.3 Clinical presentation and associated pathology
44.4 Physical exam
44.5 Imaging
44.5.1 X-rays
44.5.2 Weight-bearing CT scan
44.6 Biomechanical changes of pes cavus
44.6.1 Hindfoot
44.6.2 Forefoot
44.6.3 Soft tissues
44.7 Conservative management
44.8 Surgical management
44.8.1 Overview
44.8.2 Gastrocnemius recession/Achilles tendon lengthening
44.8.3 Plantar fascia release
44.8.4 Peroneus longus to brevis transfer
44.8.5 Posterior tibial tendon lengthening and transfer
44.8.6 Dorsiflexion osteotomy or fusion of first ray
44.8.7 Calcaneal osteotomy
44.8.8 Modified jones procedure
44.8.9 Lesser toe deformities
44.8.10 Triple arthrodesis and midfoot fusion
44.8.11 Treatment of associated ankle pathology
44.8.12 Postoperative care
44.8.13 Complications
44.9 Areas of future research
References
45 Biomechanics of Hindfoot Fusions
45.1 Introduction
45.2 Complex hindfoot biomechanics
45.3 Conditions that may require hindfoot fusion
45.3.1 Flatfoot
45.3.2 Cavus foot syndromes
45.3.3 Rheumatoid arthritis
45.3.4 Osteoarthritis
45.3.5 Calcaneal fractures
45.3.6 Talar fractures and dislocations
45.3.7 Tarsal coalitions
45.3.8 Accessory navicular
45.4 Presurgical assessment
45.4.1 Clinical exam
45.5 Imaging
45.6 Goals in treatment
45.6.1 Treatment goals in flatfoot/cavus syndromes
45.6.2 Treatment goals in rheumatoid arthritis
45.6.3 Treatment goals in osteoarthritis
45.6.4 Treatment goals in calcaneal fractures
45.6.5 Treatment goals in talar fractures and dislocations
45.7 Corrective options
45.7.1 Distraction subtalar fusion
45.7.2 Lateral column lengthening
45.7.3 Double hindfoot fusion
45.7.4 Subtalar and talonavicular fusion
45.7.5 Triple arthrodesis
45.7.6 Pantalar arthrodesis
45.7.7 General complications
45.7.8 Postoperative management requirements in hindfoot fusions
45.8 Areas of future interest
References
46 Biomechanics of Foot and Ankle Fixation
46.1 Introduction
46.2 Screws
46.3 Plates
46.4 Post and screw constructs
46.5 Nails
46.6 Beams
46.7 Areas of future research
References
47 Ankle Arthroplasty and Ankle Arthrodesis
47.1 Introduction
47.2 Brief description and history of surgical techniques
47.2.1 Ankle arthrodesis
47.2.2 Ankle arthroplasty
47.3 Biomechanical factors in presurgical assessment and consideration of arthroplasty or arthrodesis
47.3.1 Limb alignment with arthroplasty
47.3.2 Bony and ligamentous ankle anatomy with arthroplasty
47.3.3 Ankle motion in the three cardinal planes with arthroplasty
47.3.4 Bony morphology and alignment with arthrodesis
47.4 Biomechanical considerations/complications of arthroplasty or arthrodesis
47.4.1 Ankle alignment/malalignment
47.4.2 Gait mechanics
47.4.3 Arthritis at distal joints
47.4.4 Component wear or failure
47.4.5 Cadaveric gait simulation of arthroplasty and arthrodesis
47.4.6 Computational models of arthroplasty and arthrodesis
47.5 Biomechanical outcomes
47.6 Clinical outcomes
47.6.1 Safety
47.6.2 Effectiveness
47.6.3 Costs
47.6.4 Patient subgroups
47.7 Areas of future biomechanical research
References
48 Prosthetic Feet
48.1 Introduction
48.2 Prescription and expected use of prosthetic feet
48.2.1 Activity levels
48.2.2 Activity bouts and durations
48.2.3 Activity in different environments
48.3 Form of prosthetic feet
48.3.1 Solid ankle cushioned heel
48.3.2 Fixed-angle stiffness
48.3.3 Variable-angle stiffness
48.3.4 Variable stiffness
48.3.5 Powered prosthetic feet
48.4 Function of prosthetic feet
48.4.1 Clinical trials
48.4.1.1 Clinical trials examining effects of stiffness
48.4.1.2 Clinical trials examining effects of damping
48.4.1.3 Powered prosthetic feet
48.4.2 Mechanical property tests
48.4.2.1 Stiffness, hysteresis, and energy
48.4.2.2 Roll-over shape
48.4.3 Musculoskeletal modeling and simulation
48.5 Future prosthetic foot research
References
Index

Citation preview

Foot and Ankle Biomechanics

Foot and Ankle Biomechanics

Editors William R. Ledoux RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; Department of Mechanical Engineering, University of Washington, Seattle, WA, United States; Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Scott Telfer Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States; RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; Department of Mechanical Engineering, University of Washington, Seattle, WA, United States

Academic Press is an imprint of Elsevier 125 London Wall, London EC2Y 5AS, United Kingdom 525 B Street, Suite 1650, San Diego, CA 92101, United States 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom Copyright © 2023 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www. elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. ISBN: 978-0-12-815449-6 For Information on all Academic Press publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Mara E. Conner Acquisitions Editor: Carrie L. Bolger Editorial Project Manager: Clark M. Espinosa Production Project Manager: Fahmida Sultana Cover Designer: Vicky Pearson Esser Typeset by MPS Limited, Chennai, India

List of contributors David Allan Centre for Biomechanics and Rehabilitation Technologies, Staffordshire University, Stoke-on-Trent, United Kingdom Donald D. Anderson Departments of Orthopedics & Rehabilitation, Biomedical Engineering, and Industrial & Systems Engineering, University of Iowa, Iowa City, IA, United States

Sicco A. Bus Department of Rehabilitation Medicine, Amsterdam UMC, University of Amsterdam, Amsterdam Movement Sciences, Amsterdam, The Netherlands Emanuele Luigi Carniel Department of Industrial Engineering, University of Padova, Padova, Italy; Centre for Mechanics of Biological Materials, University of Padova, Padova, Italy

Anton Arndt Department of Clinical Science, Intervention and Technology, Karolinska Institute, Stockholm, Sweden; The Swedish School of Sport and Health Sciences (GIH), Stockholm, Sweden

Panagiotis Chatzistergos Centre for Biomechanics and Rehabilitation Technologies, Staffordshire University, Stoke-on-Trent, United Kingdom

John B. Arnold Allied Health & Human Performance Unit, University of South Australia, Adelaide, SA, Australia

Sagar S. Chawla Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Patrick Aubin RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; Department of Mechanical Engineering, University of Washington, Seattle, WA, United States; Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Nachiappan Chockalingam Centre for Biomechanics and Rehabilitation Technologies, Staffordshire University, Stoke-on-Trent, United Kingdom

Ruth Barn School of Health and Life Sciences, Glasgow Caledonian University, Glasgow, United Kingdom Sara Behforootan Imperial College London, Department of Surgery and Cancer, London, United Kingdom Alexander Berardo-Cates RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States Alessio Bernasconi Foot and Ankle Unit, Royal National Orthopaedic Hospital, Stanmore, Greater London, United Kingdom David Boe Rombolabs, rombolabs.github.io William Braaksma Orthopedic Associates Huron, Port Huron, MI, United States

of

Port

Claire Brockett Institute of Medical & Biological Engineering, University of Leeds, Leeds, United Kingdom Christina L. Brunnquell Department of Radiology, University of Washington, Seattle, WA, United States

Christine B. Chung UC San Diego Department of Radiology, La Jolla, CA, United States; VA San Diego Healthcare System, Department of Radiology, San Diego, CA, United States Matthew S. Conti Hospital for Special Surgery (HSS), New York, NY, United States Brian H. Dalton School of Health and Exercise Sciences, University of British Columbia Okanagan, Kelowna, BC, Canada Irene S. Davis Spaulding National Running Center, Physical Medicine and Rehabilitation, Harvard Medical School, Cambridge, MA, United States Cesar de Cesar Netto Department of Orthopedics and Rehabilitation, Carver College of Medicine, University of Iowa, Iowa City, Iowa, United States Jonathan Deland Orthopaedic Surgery, Weill Cornell Medical College, Attending, Department of Orthopedics - Foot and Ankle, HSS, New York, NY, United States Luke Donovan Department of Applied Physiology, Health, and Clinical Sciences, University of North Carolina at Charlotte, Charlotte, NC, United States

xxi

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List of contributors

Tobin Eckel Walter Reed National Military Medical Center, Bethesda, MD, United States Scott J. Ellis Hospital for Special Surgery (HSS), New York, NY, United States Tim Finkenstaedt UC San Diego Department of Radiology, La Jolla, CA, United States; Institute of Diagnostic and Interventional Radiology, University Hospital Zurich, Zurich, Switzerland; Institute of Radiology and Nuclear Medicine, Kantonsspital Winterthur, Winterthur, Switzerland Chiara Giulia Fontanella Department of Industrial Engineering, University of Padova, Padova, Italy; Centre for Mechanics of Biological Materials, University of Padova, Padova, Italy Christina E. Freibott Department of Orthopedic Surgery, Columbia University Medical Center, New York, NY, United States Jonathan H. Garfinkel Hospital for Special Surgery (HSS), New York, NY, United States Yvonne M. Golightly University of North Carolina, Chapel Hill, NC, United States Aerie Grantham RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; Department of Mechanical Engineering, University of Washington, Seattle, WA, United States Thomas M. Greiner Department of Health Professions, University of Wisconsin—La Crosse, La Crosse, WI, United States Justin K. Greisberg Department of Orthopedic Surgery, Columbia University Medical Center, New York, NY, United States

Sheree Hurn Faculty of Health, School of Clinical Sciences, Podiatry, Queensland University of Technology (QUT), Brisbane, QLD, Australia Joseph M. Iaquinto RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; Department of Mechanical Engineering, University of Washington, Seattle, WA, United States; Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States Eon Kang Interdisciplinary Consortium on Advanced Motion Performance (iCAMP), Division of Vascular Surgery and Endovascular Therapy, Michael E. DeBakey Department of Surgery, Baylor College of Medicine, Houston, TX, United States; Department of Bioengineering, University of Texas, Dallas, TX, United States Luke A. Kelly School of Human Movement and Nutrition Sciences, The University of Queensland, St Lucia, QLD, Australia

Gu

Nathan Kiewiet Orthopedic Health of Kansas City, North Kansas City, MO, United States Glenn K. Klute Center for Limb Loss and MoBility (CLiMB), Department of Veterans Affairs, Seattle, WA, United States; Department of Mechanical Engineering, University of Washington, Seattle, WA, United States Joseph J. Krzak Physical Therapy Program, Midwestern University, Downers Grove, IL, United States; Motion Analysis Center, Shriners Children’s, Chicago, IL, United States

Joseph Hamill Biomechanics Laboratory, Department of Kinesiology, University of Massachusetts Amherst, Amherst, MA, United States Marian T. Hannan Harvard Medical School, Hebrew SeniorLife, Boston, MA, United States

William R. Ledoux RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; Department of Mechanical Engineering, University of Washington, Seattle, WA, United States; Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Gordon Hendry School of Health and Life Sciences, Glasgow Caledonian University, Glasgow, United Kingdom

Morgan E. Leslie RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States

Howard J. Hillstrom Hospital for Special Surgery (HSS), New York, NY, United States

Franc¸ois Lintz Ankle and Foot Surgery Center, Clinique de l’Union, Toulouse, Occitanie, France

Rajshree Hillstrom Biomed Consulting, New York, NY, United States Bruce Elliot Hirsch Department of Neurobiology and Anatomy, Drexel University College of Medicine, Philadelphia, PA, United States

Jason T. Long Motion Analysis Lab, Cincinnati Children’s Hospital Medical Center, Cincinnati, OH, United States; Department of Orthopaedic Surgery, University of Cincinnati, College of Medicine, Cincinnati, OH, United States

Karsten Hollander Spaulding National Running Center, Physical Medicine and Rehabilitation, Harvard Medical School, Cambridge, MA, United States

Arne Lundberg Department of Clinical Science, Intervention and Technology, Karolinska Institute, Stockholm, Sweden

List of contributors

Dante Marconi Kingsbrook Jewish Medical Center, Shore Physicians Group, Somers Point, Brooklyn, NY, United States Hylton B. Menz School of Allied Health, Human Services and Sport, La Trobe University, Melbourne, VIC, Australia; La Trobe Sport and Exercise Medicine Research Centre, La Trobe University, Melbourne, VIC, Australia

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Eric Rombokas Rombolabs, rombolabs.github.io Dieter Rosenbaum Institute of Experimental Musculoskeletal Medicine, Movement Analysis Laboratory, University Hospital Muenster, Muenster, Germany

Karen J. Mickle School of Environmental and Life Sciences, University of Newcastle, NSW, Australia

Elizabeth Russell Esposito DoD-VA Extremity Trauma and Amputation Center of Excellence, Joint Base San Antonio–Fort Sam Houston, San Antonio, TX, United States; Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States

Max Mifsud Oxford University Hospitals Foundation Trust, Oxford, United Kingdom

NHS

Andrew K. Sands New York Downtown Orthopaedic Associates, New York, NY, United States

Oliver Morgan Stryker, Centennial Park, Elstree, United Kingdom

Scott Shawen OrthoCarolina Foot and Ankle Institute, Charlotte, NC, United States

Roozbeh Naemi Centre for Biomechanics and Rehabilitation Technologies, Staffordshire University, Stoke-on-Trent, United Kingdom

Palanan Siriwanarangsun UC San Diego Department of Radiology, La Jolla, CA, United States; Department of Radiology, Siriraj Hospital, Mahidol University, Bangkok Noi, Bangkok, Thailand

Bijan Najafi Interdisciplinary Consortium on Advanced Motion Performance (iCAMP), Division of Vascular Surgery and Endovascular Therapy, Michael E. DeBakey Department of Surgery, Baylor College of Medicine, Houston, TX, United States Arturo Nicola Natali Department of Industrial Engineering, University of Padova, Padova, Italy; Centre for Mechanics of Biological Materials, University of Padova, Padova, Italy Daniel C. Norvell RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; Department of Rehabilitation Medicine, University of Washington, Seattle, WA, United States Kade L. Paterson Centre for Health, Exercise and Sports Medicine, Department of Physiotherapy, School of Health Sciences, Faculty of Medicine Dentistry & Health Sciences, The University of Melbourne, Melbourne, VIC, Australia Michael T. Perez Department of Biomedical Engineering, Virginia Commonwealth University, Richmond, VA, United States Geoffrey A. Power Department of Human Health and Nutritional Sciences, College of Biological Sciences, University of Guelph, Guelph, ON, Canada Kalyani Rajopadhye Sigvard T. Hansen Foot and Ankle Institute, Physical Therapy and Hand Therapy Clinic, Harborview Medical Center, University of Washington, Seattle, WA, United States Anthony Redmond Clinical Biomechanics and Physical Medicine, University of Leeds, Leeds, United Kingdom

Michelle D. Smith School of Health & Rehabilitation Sciences, The University of Queensland, St Lucia, QLD, Australia Jinsup Song Temple University School of Podiatric Medicince, Philadelphia, PA, United States Julie Stebbins Oxford University Hospitals NHS Foundation Trust, Oxford, United Kingdom; Nuffield Department of Orthopaedics, Rheumatology and Musculoskeletal Sciences (NDORMS), Oxford University, Oxford, United Kingdom Amanda Stone RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; ARCCA, Incorporated, Seattle, WA, United States Scott Telfer Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States; RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; Department of Mechanical Engineering, University of Washington, Seattle, WA, United States Eric Thorhauer RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States; Department of Mechanical Engineering, University of Washington, Seattle, WA, United States Danielle Torp Department of Athletic Training and Clinical Nutrition, University of Kentucky, Lexington, KY, United States Robert Turner Hospital for Special Surgery (HSS), New York, NY, United States

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List of contributors

Jennifer S. Wayne Department of Biomedical Engineering, Virginia Commonwealth University, Richmond, VA, United States; Department of Biomedical Engineering & Mechanics, Virginia Tech, Blacksburg, VA, United States Gillian Weir Biomechanics Laboratory, Department of Kinesiology, University of Massachusetts Amherst, Amherst, MA, United States

Jason Wilken Department of Physical Therapy and Rehabilitation Science, University of Iowa, Iowa City, IA, United States James Woodburn Griffith Centre of Biomedical and Rehabilitation Engineering (GCORE) and School of Health Sciences and Social Work, Griffith University, QLD, Australia

Dedication This book is dedicated to Peter Cavanagh, PhD, whose career and kindness, and absolute commitment to the pursuit of excellence, continue to serve as an inspiration to biomechanists around the world.

Contents List of contributors Preface Anatomical Terms Used in Foot and Ankle Biomechanics

xxi xxv xxvii

Part 1 Introduction 1. Anatomy of the Foot

3

Bruce Elliot Hirsch 1.1 Skeletal structures 1.1.1 Tibia and fibula 1.1.2 Segments of the foot skeleton 1.1.3 Talus 1.1.4 Calcaneus 1.1.5 Navicular 1.1.6 Cuboid 1.1.7 Medial cuneiform 1.1.8 Intermediate cuneiform 1.1.9 Lateral cuneiform 1.1.10 Metatarsals 1.1.11 Phalanges 1.2 Joints 1.2.1 Tibiofibular syndesmosis 1.2.2 Ankle joint 1.2.3 Subtalar joint 1.2.4 Talocalcaneonavicular joint 1.2.5 Calcaneocuboid joint 1.2.6 Cuboideonavicular joint 1.2.7 Cuneonavicular joints 1.2.8 Intercuneiform and cuneocuboid joints 1.2.9 Tarsometatarsal joints 1.2.10 Proximal intermetatarsal joints 1.2.11 Distal intermetatarsal joints 1.2.12 Lesser metatarsophalangeal joints 1.2.13 Hallucal metatarsophalangeal joint 1.2.14 Interphalangeal joints 1.3 Muscles and fascial specializations 1.3.1 Fascial specializations 1.3.2 Extrinsic dorsal muscles

3 3 3 6 6 7 7 7 7 8 8 10 11 11 11 13 14 14 17 17 17 18 18 19 19 19 20 20 21 21

1.3.3 Extrinsic plantar muscles 1.3.4 Extrinsic lateral muscles 1.3.5 Intrinsic dorsal foot muscles 1.3.6 Intrinsic plantar muscles 1.4 Nerves 1.4.1 The fibular nerves in the foot 1.4.2 The tibial nerves in the foot 1.5 Blood supply 1.5.1 Arteries 1.5.2 Veins 1.5.3 Lymphatics Further reading

2. Basic Biomechanics

25 28 29 31 38 38 40 41 41 42 43 44

45

Joseph M. Iaquinto 2.1 Introduction 2.2 Terminology 2.3 Statics 2.4 Dynamics 2.5 Strength of materials and deformation 2.6 Viscoelasticity 2.7 Summary References

45 45 46 48 48 49 50 51

3. Anatomical Nomenclature: Conundrums of Nonstandardized Foot and Ankle Terminology

53

Thomas M. Greiner 3.1 Introduction 3.2 Anatomical descriptions 3.2.1 The persistence of eponyms 3.2.2 Regional descriptions 3.2.3 Applying the anatomical position 3.2.4 Application specific human anatomical positions 3.2.5 Defining anatomical directions, planes, and axes 3.3 Foot motions 3.3.1 Defining motions 3.3.2 Whole foot motions and their complexity

53 54 54 54 55 56 61 64 64 67 vii

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Contents

3.4 Terminological implications of mathematical choices 3.5 Conclusion: standardizing foot and ankle terminology References

7. Ligaments 71 71 73

Part 2 Function 4. Kinematics and Kinetics of the Foot and Ankle during Gait

77

Jason T. Long and Joseph J. Krzak 4.1 Introduction 4.2 Overview of relevant anatomy 4.3 Overview of kinematic and kinetic modeling 4.4 Healthy and impaired feet 4.5 Multisegment foot models 4.6 Future areas of research 4.6.1 Biplane fluoroscopy 4.6.2 Modeling 4.7 Conclusion References

77 78 79 81 81 84 84 85 85 86

5. Bone, Cartilage, and Joint Function

89

Michael T. Perez and Jennifer S. Wayne 5.1 Bone components and structure 5.2 Cartilage 5.3 Joint functions 5.3.1 Talocrural joint 5.3.2 Talocalcaneal (subtalar) joint 5.3.3 Transverse tarsal joint 5.3.4 Tarsometatarsal joint 5.3.5 Metatarsophalangeal joint 5.4 Areas of future research References

6. Muscles and Tendons

89 92 95 95 96 97 99 99 100 100

103

Karen J. Mickle 6.1 Introduction 6.2 Biomechanical function 6.2.1 Normal foot 6.2.2 Aging 6.2.3 Pathologies 6.2.4 Footwear and orthoses 6.2.5 Interventions 6.3 Areas of future biomechanical research References

103 104 104 108 109 113 114 116 117

121

Aerie Grantham, Joseph M. Iaquinto and Alexander Berardo-Cates 7.1 Introduction 7.2 Ligament anatomy 7.3 Mechanical properties 7.3.1 Ankle joint 7.3.2 Hindfoot ligaments 7.3.3 Midfoot ligaments 7.3.4 Forefoot 7.3.5 Variations in mechanical properties 7.4 Ligament sprains 7.5 Overcoming limitations 7.6 Future areas of research References

8. Plantar Soft Tissue

121 121 122 122 124 124 126 126 127 127 132 132

135

William R. Ledoux 8.1 Introduction 8.2 Anatomy 8.2.1 Gross anatomy 8.2.2 Histological or biochemical 8.2.3 Medical imaging of tissue thickness 8.3 Biomechanical function 8.4 Mechanical properties 8.4.1 Structural in vivo testing 8.4.2 Structural ex vivo testing 8.4.3 Material ex vivo testing 8.4.4 Ultrasound 8.4.5 Other in vivo techniques 8.5 Effect of aging 8.6 Diabetic plantar soft tissue 8.6.1 Other pathologies associated with the plantar soft tissue 8.7 Areas of future biomechanical research References

135 135 135 136 137 140 140 140 140 142 142 143 143 144 144 145 145

Part 3 Measurement and Analysis Techniques: Kinematics and Kinetics 9. Multisegment Foot Models

151

Amanda Stone 9.1 Basic principles of multisegment foot models 9.1.1 Overview of motion capture 9.1.2 Why use a multisegment foot model?

151 151 152

Contents

9.2 Selecting an appropriate multisegment foot model 9.2.1 Segments and bone groupings 9.2.2 Marker type and placement 9.2.3 Coordinate systems 9.2.4 Offsets 9.2.5 Standardized description of multisegment foot models 9.3 Review of current multisegment foot models 9.3.1 Milwaukee kinematic model 9.3.2 Leardini/Rizzoli kinematic model 9.3.3 Oxford kinematic model 9.3.4 MacWilliams/Kinfoot kinetic model 9.3.5 Bruening kinetic model 9.3.6 Direct comparison of current multisegment foot models 9.4 Applications, considerations, and limitations 9.4.1 Clinical applications 9.4.2 Sources of error 9.5 Areas of future biomechanical research References

10. Invasive Techniques for Studying Foot and Ankle Kinematics

152 152 153 153 153 154 154 154 155 156 157 157 157 159 159 160 161 161

167

Arne Lundberg and Anton Arndt 10.1 Introduction 10.2 Early invasive studies of foot and ankle biomechanics 10.3 Radiostereometric analysis 10.3.1 Technique 10.3.2 Talocrural joint 10.3.3 Relationship between the joints distal to the talus 10.3.4 Transferral of rotation between the leg and the foot 10.3.5 Ankle mortise width 10.4 Applications and significance of studies using intracortical pins for foot and ankle kinematics 10.4.1 Skin movement artifact in foot and ankle kinematics 10.4.2 Ankle kinematics 10.4.3 Foot and ankle basic research in walking and running kinematics 10.4.4 Applied studies of orthoses and shoe conditions 10.5 Limitations and future directions

167 168 169 169 169 170

170 171

171 171 173

173 173 174

Appendix: Insertion of markers in bones of the foot and ankle References

11. Biplane Fluoroscopy

ix

175 177

179

Eric Thorhauer and William R. Ledoux 11.1 Introduction 11.2 Background and history of biplane fluoroscopy 11.2.1 Overview of how X-ray imaging works 11.2.2 Overview of biplane system history and evolution 11.3 Other techniques for tracking foot bone kinematics 11.4 Challenges specific to foot and ankle tracking with biplane fluoroscopy 11.5 Overview of biplane hardware 11.6 Overview of biplane software 11.7 Clinical biplane foot and ankle studies 11.7.1 Biplane systems consisting of two C-arms 11.7.2 Biplane systems consisting of two modified C-arms 11.7.3 Biplane systems consisting of independent X-ray sources and image intensifiers 11.8 Future applications and directions References

12. Plantar Pressure and Ground Reaction Forces

179 180 180 180 181 182 182 185 187 187 189

189 190 191

197

Dieter Rosenbaum and Scott Telfer 12.1 Introduction: clinical relevance of force and pressure measurements in foot and ankle biomechanics 12.2 Background: force versus pressure 12.2.1 History 12.2.2 Development of measurement technologies 12.2.3 Visualization and analytical options 12.3 Research applications and selected clinical examples 12.3.1 Diabetes 12.3.2 Children’s flatfoot 12.3.3 Sports 12.4 Areas of future research References

197 198 198 199 201 204 205 206 206 207 208

x

Contents

13. Electromyography and Dynamometry for Investigating the Neuromuscular Control of the Foot and Ankle 211 Brian H. Dalton and Geoffrey A. Power 13.1 Introduction 13.2 Electromyography 13.2.1 Surface electromyography 13.2.2 Indwelling electromyography and motor unit recordings 13.3 Dynamometry 13.3.1 Isometric 13.3.2 Dynamic 13.4 Ankle and foot related considerations and insights 13.4.1 Motor unit behavior and quantity 13.4.2 Maximal voluntary contractions and knee angle 13.4.3 History-dependence of force 13.5 Future research References

14. From Impossible to Unnoticed: Wearable Technologies and the Miniaturization of Grand Science

211 212 212 213 214 214 218 219 219 221 221 224 224

229

Eric Rombokas and David Boe 14.1 Introduction 14.1.1 Tech affords understanding 14.1.2 Wearable domains 14.1.3 Breakout box: what makes a successful wearable device? 14.2 The past 14.2.1 Force and pressure 14.2.2 Health and activity sensing 14.3 The present 14.3.1 Force and pressure sensing 14.3.2 Health and activity monitoring 14.3.3 Actuation and assistance 14.3.4 Haptics 14.3.5 Motion capture 14.4 The future References

229 229 230 230 231 231 232 234 234 235 236 236 237 238 239

Part 4 Measurement and Analysis Techniques: Imaging 15. Integrated Laboratories for Pursuing Pedal Pathologies

245

Oliver Morgan, Rajshree Hillstrom, Jinsup Song, Robert Turner, Marian T. Hannan,

Yvonne M. Golightly, Scott J. Ellis, Jonathan Deland and Howard J. Hillstrom 15.1 Introduction 15.2 Our method of approach 15.3 Integrated laboratories 15.3.1 Epidemiology 15.3.2 In vivo experimentation 15.3.3 In vitro experimentation 15.3.4 In silico simulation 15.4 Case study of the integrated laboratories concept to the study of hallux rigidus 15.4.1 Epidemiology 15.4.2 In vivo experimentation 15.4.3 In vitro experimentation 15.4.4 In silico simulation 15.5 Future biomechanical research References

16. Radiographs

245 246 247 247 247 248 249 253 253 254 256 258 260 262

265

Morgan Leslie and William R. Ledoux 16.1 Introduction 16.2 Radiographic technology 16.3 Standard radiographic views of the foot and ankle 16.4 Definitions of X-ray measurements of foot shape 16.5 Foot-specific applications and considerations 16.6 Clinical X-ray measures of foot shape 16.7 Issues with X-ray measures of foot shape 16.8 Areas of future biomechanical research References

17. Computed Tomography of the Foot and Ankle

265 265 267 269 270 270 272 273 274

277

Scott Telfer, Christina L. Brunnquell and William R. Ledoux 17.1 Introduction 17.1.1 History and development of computed tomography 17.1.2 Comparison to other imaging modalities 17.1.3 Computed tomography protocols for the foot and ankle 17.2 Foot-specific applications and considerations 17.2.1 Disease diagnosis 17.2.2 Surgical assessment and planning 17.2.3 Biomechanics research 17.3 Areas of future biomechanical research References

277 278 278 279 281 281 281 282 284 285

Contents

18. Weight-bearing Computed Tomography of the Foot and Ankle 289 Franc¸ois Lintz, Alessio Bernasconi and Cesar de Cesar Netto 18.1 18.2 18.3 18.4 18.5 18.6

Introduction Biases of conventional radiography Technical aspects Indications 3D biometrics Advantages and limitations of weight-bearing computed tomography 18.7 Future areas of research 18.8 Conclusion Acknowledgments Conflict of interest statement References Further reading

19. Magnetic Resonance Imaging of the Foot and Ankle

289 290 291 294 294

296 296 297 297 297 297 299

301

Tim Finkenstaedt, Palanan Siriwanarangsun and Christine B. Chung 19.1 Introduction 19.2 Magnetic resonance imaging sequences 19.3 Magnetic resonance imaging versus computed tomography 19.4 Magnetic resonance imaging appearance of musculoskeletal tissue—normal and pathology 19.4.1 Short T2 tissues 19.4.2 Tendons 19.4.3 Pseudo tendon pathology—the magic angle effect or phenomenon 19.4.4 Ligaments 19.4.5 Bone 19.4.6 Hyaline and fibrocartilaginous cartilage 19.4.7 Muscle 19.4.8 Bursa/synovia 19.5 Tailored magnetic resonance imaging protocol for the foot and ankle— indication driven 19.5.1 Imaging orientation 19.5.2 Tailored magnetic resonance imaging protocols 19.5.3 Optimized imaging planes 19.5.4 Metal artifact reduction sequences

301 302 303

304 304 304

305 305 305 307 307 307

308 308 308 310 311

19.6 Magnetic resonance imaging anatomy of the foot and ankle 19.6.1 Ankle 19.6.2 Hindfoot 19.6.3 Midfoot 19.6.4 Forefoot 19.7 Areas of future research 19.7.1 Radiation-free bone imaging 19.7.2 Magnetic resonance imaging scan time reduction 19.7.3 Volumetric, isotropic 3D magnetic resonance imaging sequences References

20. Biomechanical Assessment of Soft Tissues in the Foot and Ankle Using Ultrasound

xi

311 311 313 315 317 318 318 319

320 320

323

Roozbeh Naemi, David Allan, Sara Behforootan, Panagiotis Chatzistergos and Nachiappan Chockalingam 20.1 Introduction 323 20.1.1 Ultrasound imaging, how it works 323 20.2 Ultrasound assessment of structural changes: the effect of weightbearing activities 325 20.2.1 Assessment of plantar soft tissue thickness in relation to weightbearing activities 325 20.2.2 Assessment of plantar fascia thickness in relation to weightbearing activities 326 20.2.3 Assessment of Achilles tendon thickness in relation to weightbearing activities 327 20.2.4 Summary and limitations of weightbearing ultrasound 327 20.3 Ultrasound assessment combined with measurement of load 328 20.3.1 Ultrasound combined with load cells to assess the mechanical properties of the plantar soft tissue 328 20.3.2 Ultrasound combined with dynamometry to assess the mechanical properties of Achilles tendon 329 20.3.3 Summary and limitations of ultrasound assessment combined with measurement of load 331 20.4 Ultrasound elastography (sonoelastography) 331

xii

Contents

20.4.1 Ultrasound elastography to assess the mechanical properties of plantar soft tissue 20.4.2 Ultrasound elastography to assess the mechanical properties of plantar fascia 20.4.3 Ultrasound elastography to assess the mechanical properties of Achilles tendon 20.4.4 Summary and limitations of ultrasound elastography 20.5 Conclusion and future areas of research References

21. 3D Surface Scanning of the Foot and Ankle

23. Finite Element Modeling 331

333

333 334 335 335

339

Scott Telfer 21.1 Introduction 21.1.1 The history and development of 3D scanning 21.1.2 Technologies 21.2 Foot-specific applications and considerations 21.2.1 Reliability and comparisons to other techniques 21.2.2 Population studies 21.2.3 Orthoses and footwear 21.2.4 Musculoskeletal models 21.2.5 Bones and other internal anatomy 21.3 Areas of future biomechanical research References

339 340 341 342 342 343 344 344 344 345 345

Part 5 Measurement and Analysis Techniques: Simulation and Modeling 22. Cadaveric Gait Simulation

351

365

Panagiotis Chatzistergos, Sara Behforootan, Roozbeh Naemi and Nachiappan Chockalingam 23.1 Introduction 23.2 Basic concepts of finite element modeling 23.3 Applications of finite element analysis in foot biomechanics 23.3.1 Simulation of the interaction between foot and footwear 23.3.2 Simulation of healthy foot biomechanics 23.3.3 Finite element modeling for the in vivo material characterization of soft tissues 23.3.4 Simulation of pathological conditions 23.4 Modeling strategies 23.4.1 Geometry design 23.4.2 Meshing 23.4.3 Material properties 23.4.4 Solver selection 23.4.5 Reliability assessment 23.5 Limitations and future research toward clinically applicable finite element modeling 23.6 Summary References

24. Musculoskeletal Modeling of the Foot and Ankle

365 365 366 366 369

371 373 375 375 379 380 381 382

383 383 384

387

Scott Telfer 24.1 Introduction 24.1.1 Musculoskeletal models and their development 24.1.2 Validation techniques 24.1.3 Challenges in modeling the foot and ankle 24.2 Foot specific models and applications 24.3 Areas of future biomechanical research References

387 388 390 391 392 394 395

Patrick Aubin and William R. Ledoux 22.1 22.2 22.3 22.4

Introduction Techniques for dynamic gait simulation Limitations of dynamic gait simulation Clinical applications of dynamic gait simulation 22.5 Areas of future biomechanical research 22.6 Conclusion References

351 352 354 359 360 360 361

25. Predicting and Preventing Posttraumatic Osteoarthritis of the Ankle

397

Donald D. Anderson, Jason Wilken, Claire Brockett and Anthony Redmond 25.1 Introduction: pathomechanical origins of posttraumatic osteoarthritis

397

Contents

25.2 Pathomechanics I: acute joint injury severity 25.3 Pathomechanics II: chronic stress aberration 25.4 Pathomechanics III: altered kinematics 25.5 Areas of future biomechanical research 25.5.1 Posttraumatic ankle osteoarthritis: opportunities for intervention informed by pathomechanical knowledge 25.6 Summary/conclusion References

26. Mechanics of Biological Tissues

397 399 402 404

404 407 407

411

Arturo Nicola Natali, Emanuele Luigi Carniel and Chiara Giulia Fontanella 26.1 Introduction 26.2 Materials and methods 26.2.1 Finite element modeling of the foot 26.2.2 Formulation of constitutive models 26.2.3 Identification of constitutive parameters 26.2.4 Numerical analyses of foot functionality 26.2.5 Limitations of computational modeling 26.2.6 Future biomechanics research 26.3 Conclusion References

411 412 412 412 417 421 425 426 426 427

Part 6 Clinical Biomechanics of the Foot and Ankle 27. Clinical Examination of the Foot and Ankle Introduction Demographics Vital signs Patient history Assessment of pain Visual observation/inspection 27.6.1 Skin 27.6.2 Edema 27.6.3 Atrophy 27.6.4 Temperature

28. Foot Type Biomechanics

436 436 437 438 438 438 439 439 440 440 440 441 442 442 443 443 444 444 444 445 445 446 446

451

Scott Telfer and William R. Ledoux 28.1 28.2 28.3 28.4 28.5

Introduction Structural foot type Functional foot type Foot type biomechanics Association with pain and injury 28.5.1 Pain and injury 28.6 Treatments 28.7 Areas of future biomechanical research References

451 452 454 454 456 456 457

29. Traumatic Foot and Ankle Injuries

461

457 458

433

Kalyani Rajopadhye 27.1 27.2 27.3 27.4 27.5 27.6

27.6.5 Scarring 27.6.6 Callus patterns 27.6.7 Exostoses 27.6.8 Ankle and foot deformities 27.7 Lower extremity alignment 27.8 Foot posture or foot shape 27.8.1 Planus foot type 27.8.2 Cavus foot type 27.9 Limb length 27.10 Radiographic examination 27.11 Range of motion/flexibility/joint mobility 27.12 Joint mobility 27.13 Ligamentous/stability testing 27.14 Tendon 27.15 Muscle strength 27.16 Sensory testing 27.17 Circulation 27.18 Foot and ankle specific testing 27.19 Footwear examination 27.20 Functional assessment 27.21 Outcomes assessment 27.22 Areas of future biomechanical research References

xiii

433 433 433 433 434 434 434 435 435 435

Scott Shawen and Tobin Eckel 29.1 Introduction 29.2 Pilon fractures 29.2.1 Etiology and pathophysiology 29.2.2 Symptoms 29.2.3 Diagnostics/classification 29.2.4 Treatment 29.3 Calcaneal fractures 29.3.1 Etiology and pathophysiology

461 461 461 461 461 462 464 464

xiv

Contents

29.3.2 Symptoms 29.3.3 Diagnostics/classification 29.3.4 Treatment 29.4 Talus fractures 29.4.1 Etiology and pathophysiology 29.4.2 Symptoms 29.4.3 Diagnostics/classification 29.4.4 Treatment 29.5 Tarsometatarsal (Lisfranc) injuries 29.5.1 Etiology and pathophysiology 29.5.2 Symptoms 29.5.3 Diagnostics/classification 29.5.4 Treatment 29.6 Metatarsal fractures 29.6.1 Etiology and pathophysiology 29.6.2 Symptoms 29.6.3 Diagnostics/classification 29.6.4 Treatment 29.7 Midfoot crush injuries 29.7.1 Etiology and pathophysiology 29.7.2 Symptoms 29.7.3 Treatment 29.8 Acute ankle sprains 29.8.1 Etiology and pathophysiology 29.8.2 Symptoms 29.8.3 Treatment 29.9 Syndesmosis tears 29.9.1 Etiology and pathophysiology 29.9.2 Symptoms 29.9.3 Treatment 29.10 Achilles tendon rupture 29.10.1 Etiology and pathophysiology 29.10.2 Symptoms 29.10.3 Treatment 29.11 Areas of future research References

30. The Pediatric Foot

464 464 464 465 465 465 465 466 467 467 468 468 468 468 468 469 469 470 472 472 472 472 473 473 473 473 473 473 473 473 474 474 474 474 475 475

477

Julie Stebbins and Max Mifsud 30.1 Introduction 30.2 Common pathologies affecting pediatric feet 30.2.1 Congenital foot deformities 30.2.2 Developmental foot deformities 30.2.3 Foot pathologies associated with other conditions 30.3 Functional assessment of the pediatric foot 30.4 Areas for future research References

477 478 479 482 482 484 485 486

31. Neurological Foot Pathology

489

Morgan E. Leslie and Joseph M. Iaquinto 31.1 Introduction 31.2 Stroke 31.2.1 Pathology related to the musculoskeletal system 31.2.2 Impact on kinematics 31.2.3 Impact on foot function 31.2.4 Clinical treatment 31.3 Cerebral palsy 31.3.1 Definition 31.3.2 Structural deformities and gait deviations 31.3.3 Treatment 31.4 Toe walking 31.4.1 Diagnosis and etiology 31.4.2 Biomechanical and musculoskeletal function 31.4.3 Treatment 31.5 Peripheral neuropathy 31.5.1 Background 31.5.2 Musculoskeletal and movement implications 31.6 Foot drop 31.6.1 Pathology 31.6.2 Impact on biomechanics 31.6.3 Treatment 31.7 Tarsal tunnel syndrome 31.7.1 Pathology 31.7.2 Impact on biomechanics 31.7.3 Treatment 31.8 Morton’s neuroma 31.8.1 Pathology 31.8.2 Impact on biomechanics 31.8.3 Treatment 31.9 Charcot foot 31.9.1 Pathology 31.9.2 Impact on biomechanics 31.9.3 Treatment 31.10 Charcot-Marie-Tooth disease 31.10.1 Pathology 31.10.2 Impact on biomechanics— pediatric and young adult 31.10.3 Impact on biomechanics 31.10.4 Treatment 31.11 Friedreich’s ataxia 31.11.1 Background and pathology 31.11.2 Gait analysis 31.11.3 Musculoskeletal effects 31.11.4 Clinical treatment 31.12 Poliomyelitis 31.12.1 Pathology

489 489 489 490 491 492 492 492 493 494 494 494 494 495 495 495 496 496 496 497 497 498 498 498 498 499 499 499 499 499 499 500 500 500 500

501 501 501 502 502 502 502 502 503 503

Contents

31.12.2 Current status 31.12.3 Impact on biomechanics 31.12.4 Treatment 31.13 Areas of Future Research References

503 503 504 504 504

32. Chronic Foot and Ankle Injuries

507

Danielle Torp and Luke Donovan 32.1 Introduction 32.1.1 Chronic injury through microtrauma 32.1.2 Chronic injury through macrotrauma 32.1.3 Impairment-based rehabilitation model for treating chronic injuries 32.1.4 Role of patient-oriented outcomes 32.2 Chronic ankle instability 32.2.1 Anatomy overview 32.2.2 Etiology 32.2.3 Clinical impairments 32.2.4 Treatment 32.3 Plantar fasciitis 32.3.1 Anatomical overview 32.3.2 Etiology 32.3.3 Clinical impairments 32.3.4 Treatment 32.4 Tendinopathy (Achilles, peroneal, and posterior tibialis) 32.4.1 Anatomical overview 32.4.2 Etiology 32.4.3 Clinical impairments 32.4.4 Treatment 32.5 Stress fractures (navicular, metatarsals) 32.5.1 Anatomical overview 32.5.2 Etiology 32.5.3 Clinical impairments 32.5.4 Treatment 32.6 Sesamoiditis 32.6.1 Anatomical overview 32.6.2 Etiology 32.6.3 Clinical impairments 32.6.4 Treatment 32.7 Retrocalcaneal bursitis 32.7.1 Anatomical overview 32.7.2 Etiology 32.7.3 Clinical impairments 32.7.4 Treatment 32.8 Areas of future research for chronic foot and ankle injuries References

507 507 509

509 510 510 510 510 512 512 513 513 513 514 514 514 515 515 516 516 516 516 516 517 518 518 518 518 518 520 520 520 520 520 520 520 523

33. Hallux Valgus

xv

527

Sheree Hurn 33.1 Introduction 33.2 Prevalence 33.3 Etiology 33.3.1 Genetics and race 33.3.2 Structural and biomechanical factors 33.3.3 Footwear 33.4 Diagnosis and imaging 33.4.1 Clinical diagnosis 33.4.2 Radiographic assessment 33.4.3 Ultrasound 33.4.4 Computed tomography 33.4.5 Magnetic resonance imaging 33.5 Clinical presentation 33.5.1 Foot pain 33.5.2 Footwear 33.5.3 Self-reported function and quality of life 33.6 Functional outcomes 33.6.1 Balance and falls 33.6.2 Hallux flexion and abduction 33.6.3 Gait analysis 33.7 Treatment pathways 33.7.1 Nonsurgical treatment 33.7.2 Surgical treatment 33.8 Future directions for research 33.9 Summary References

34. Osteoarthritis of the Foot and Ankle

527 527 528 528 528 529 529 529 530 532 533 533 534 534 534 534 534 534 535 535 536 536 540 541 541 541

547

Kade L. Paterson, Luke A. Kelly and Michelle D. Smith 34.1 Introduction 34.1.1 Osteoarthritis symptoms and diagnosis 34.1.2 Structural changes 34.1.3 Risk factors and classification of osteoarthritis 34.1.4 Foot and ankle osteoarthritis subtypes 34.2 First metatarsophalangeal joint osteoarthritis 34.2.1 Etiology and impact 34.2.2 Clinical findings 34.2.3 Structural and biomechanical features 34.2.4 Clinical and biomechanical effects of conservative treatment

547 547 548 549 549 549 549 549 550 551

xvi

Contents

34.3 Midfoot osteoarthritis 34.3.1 Etiology and impact 34.3.2 Clinical findings 34.3.3 Structural and biomechanical features 34.3.4 Clinical and biomechanical effects of conservative treatment 34.4 Ankle osteoarthritis 34.4.1 Etiology and impact 34.4.2 Clinical findings 34.4.3 Structural and biomechanical features 34.4.4 Clinical and biomechanical effects of conservative treatment 34.5 Areas of future biomechanical research 34.6 Summary References

35. Diabetic Foot Disease

552 552 553 553

555 556 556 556 557

557

581

James Woodburn, Ruth Barn and Gordon Hendry 36.1 Introduction 36.2 Rheumatoid arthritis 36.2.1 Early rheumatoid arthritis 36.2.2 Established rheumatoid arthritis 36.3 Spondlyarthropathies 36.4 Juvenile idiopathic arthritis 36.5 Connective tissue disorders 36.6 Gout 36.7 Future research References

37. The Aging Foot

581 582 582 582 589 590 591 591 592 592

595

John B. Arnold and Hylton B. Menz 558 559 560

565

Bijan Najafi and Gu Eon Kang 35.1 Background on diabetes 35.2 Overview of key negative outcomes of diabetic foot disease 35.2.1 Diabetic peripheral neuropathy 35.2.2 Peripheral vascular disease 35.2.3 Diabetic plantar ulceration 35.2.4 Foot deformities 35.2.5 Lower-extremity fractures 35.2.6 Charcot neuroarthropathy 35.2.7 Lower extremity amputation 35.3 Risk factors for the development and progression of diabetic foot disease 35.4 Changes in kinematics and kinetics in diabetic foot disease 35.5 Changes in tissue characteristics 35.5.1 Muscle: fatty infiltration and reduction of intrinsic foot muscle volumes 35.5.2 Bone 35.5.3 Cartilage 35.5.4 Tendon 35.5.5 Plantar fascia 35.6 The relationship between foot deformities and plantar ulceration 35.7 The relationship between lower extremity fractures and Charcot neuropathic osteoarthropathy 35.8 Areas of future biomechanical research References

36. Rheumatic Foot Disease

565 566 566 567 567 568 568 568 570 571 571 573

573 573 573 574 574 574

574 575 575

37.1 Changing properties and functions of foot tissues 595 37.1.1 Bone 595 37.1.2 Cartilage 596 37.1.3 Muscle 596 37.1.4 Tendon 597 37.1.5 Ligament 598 37.1.6 Skin 598 37.1.7 Neural 598 37.1.8 Fat pad 599 37.2 Foot posture and morphology 599 37.2.1 Anthropometrics 599 37.2.2 Foot posture 599 37.2.3 Arch height 600 37.2.4 Joint range-of-motion 600 37.3 Foot function (kinematics/kinetics/ plantar pressures) 601 37.3.1 Kinetics 601 37.3.2 Kinematics 603 37.4 Foot posture, foot disorders, and mobility limitations 603 37.4.1 Foot posture and foot deformity 603 37.4.2 Foot posture and foot symptoms 603 37.4.3 Foot posture, mobility limitations, and falls 604 37.5 Areas for future research 604 References 605

38. Biomechanics of Athletic Footwear 611 Gillian Weir and Joseph Hamill 38.1 Introduction 38.2 Anatomy of a running shoe 38.3 Biomechanics of athletic footwear design 38.3.1 Cushioning 38.3.2 Hindfoot stability

611 612 612 613 614

Contents

38.4 Types of shoes and their features 38.4.1 Casual shoes 38.4.2 Running shoes 38.4.3 Racing flats & spikes 38.4.4 Marathon shoes 38.4.5 Other sports shoes 38.4.6 New shoe innovations 38.4.7 Graphene outsoles 38.5 Shod versus barefoot 38.6 Footwear related injuries 38.7 Future footwear research References

39. Minimal Shoes: Restoring Natural Running Mechanics

615 615 615 615 616 616 617 618 618 619 620 620

623

40.2.2 Kinetic effects of foot orthosis 40.2.3 Effects of foot orthosis on plantar pressure 40.2.4 Effects of foot orthosis on muscle activity patterns 40.3 Effects of foot orthosis on clinical conditions 40.3.1 Rheumatoid arthritis 40.3.2 Symptomatic flat foot 40.3.3 Heel pain (plantar fasciitis) 40.3.4 Osteoarthritis 40.3.5 Sports injuries and other conditions 40.4 Areas of future research References

xvii

641 641 642 642 642 643 643 643 643 643 644

Karsten Hollander and Irene S. Davis 39.1 Introduction 39.2 Brief history of running footwear 39.3 Biomechanics of barefoot and conventional shod running 39.3.1 Barefoot running pattern 39.3.2 Conventional shod running pattern 39.3.3 Comparison of mechanics between conventional shod and barefoot running 39.4 Minimal footwear running 39.4.1 Definition of full minimalist footwear 39.4.2 Comparison of full minimal to barefoot running 39.4.3 Comparison of full minimal to partial minimal shoes 39.4.4 Comparison of full minimal to conventional footwear 39.5 Effect of minimal shoes on the foot musculoskeletal system 39.6 Summary 39.7 Future research References

623 623 626 626 626

626 628 628 629 629 630 630 631 631 632

Part 7 Clincial Interventions 40. Foot Orthoses

637

41. Ankle-Foot Orthoses and Rocker Bottom Shoes

647

Elizabeth Russell Esposito 41.1 Introduction 41.2 Ankle-foot orthoses 41.2.1 Controlling rotational motion 41.2.2 Controlling translational motion 41.2.3 Controlling axial forces 41.2.4 Altering the line of action of the ground reaction force 41.3 Rocker bottom shoes 41.4 Roll-over shape 41.5 Patient populations 41.5.1 Stroke 41.5.2 Cerebral palsy 41.5.3 Ankle arthritis 41.5.4 Limb salvage 41.6 Design and prescription of ankle-foot orthosis 41.6.1 Conventional versus advanced 41.6.2 Articulated versus nonarticulated 41.7 Design and prescription of rocker bottom shoes 41.8 Variations on materials 41.9 New designs 41.10 Sport applications 41.11 Areas of future research References

647 647 648 648 648 648 648 649 649 650 651 651 652 652 652 653 654 655 655 655 656 656

Scott Telfer 40.1 Introduction 40.1.1 Design and manufacture of foot orthoses 40.2 Biomechanical effects of foot orthoses 40.2.1 Kinematic effects of foot orthosis

637 637 639 640

42. Diabetic Footwear

661

Sicco A. Bus 42.1 Introduction 42.2 Foot biomechanics and offloading

661 661

xviii

Contents

42.3 The biomechanical effect of diabetic footwear and offloading devices 42.4 Footwear and offloading for ulcer healing 42.5 Diabetic footwear for ulcer prevention 42.6 Footwear and offloading adherence 42.7 Other considerations 42.8 Future research 42.9 Conclusions References

662 663 664 665 666 666 666 667

43. Reconstructions for Adult-acquired Flatfoot Deformity 669 Matthew S. Conti, Jonathan H. Garfinkel and Scott J. Ellis 43.1 Introduction 43.2 Hindfoot valgus 43.2.1 Bony anatomy 43.2.2 Ligament failure 43.2.3 Surgical reconstruction 43.3 Forefoot external rotation 43.3.1 Bony anatomy 43.3.2 Ligament and tendon failure 43.3.3 Surgical reconstruction 43.4 Sag at the talonavicular joint 43.4.1 Bony anatomy 43.4.2 Ligamentous failure 43.4.3 Surgical reconstruction 43.5 Failure of the posterior tibial tendon 43.5.1 Clinical assessment 43.5.2 MRI assessment 43.5.3 Surgical reconstruction 43.6 Gastrocnemius and Achilles tightness 43.6.1 Clinical assessment 43.6.2 Surgical treatment 43.7 Medial arch eversion 43.7.1 Clinical assessment 43.7.2 Bony anatomy 43.7.3 Ligamentous failure 43.7.4 Surgical reconstruction 43.8 Special considerations 43.9 Future biomechanical studies and conclusion References

44. Cavus Foot Reconstructions

669 670 670 671 672 674 674 675 675 677 677 678 678 679 680 680 680 680 681 681 681 681 682 682 682 682 683 683

687

Nathan Kiewiet and William Braaksma 44.1 Introduction 44.2 Etiology 44.2.1 Traumatic 44.2.2 Neurologic

687 687 687 687

44.2.3 Residual clubfoot 44.2.4 Idiopathic 44.3 Clinical presentation and associated pathology 44.4 Physical exam 44.5 Imaging 44.5.1 X-rays 44.5.2 Weight-bearing CT scan 44.6 Biomechanical changes of pes cavus 44.6.1 Hindfoot 44.6.2 Forefoot 44.6.3 Soft tissues 44.7 Conservative management 44.8 Surgical management 44.8.1 Overview 44.8.2 Gastrocnemius recession/ Achilles tendon lengthening 44.8.3 Plantar fascia release 44.8.4 Peroneus longus to brevis transfer 44.8.5 Posterior tibial tendon lengthening and transfer 44.8.6 Dorsiflexion osteotomy or fusion of first ray 44.8.7 Calcaneal osteotomy 44.8.8 Modified jones procedure 44.8.9 Lesser toe deformities 44.8.10 Triple arthrodesis and midfoot fusion 44.8.11 Treatment of associated ankle pathology 44.8.12 Postoperative care 44.8.13 Complications 44.9 Areas of future research References

45. Biomechanics of Hindfoot Fusions

688 688 688 689 690 690 691 691 691 692 692 692 693 693 693 694 694 694 694 695 696 696 697 698 698 699 699 699

701

Dante Marconi and Andrew K. Sands 45.1 Introduction 45.2 Complex hindfoot biomechanics 45.3 Conditions that may require hindfoot fusion 45.3.1 Flatfoot 45.3.2 Cavus foot syndromes 45.3.3 Rheumatoid arthritis 45.3.4 Osteoarthritis 45.3.5 Calcaneal fractures 45.3.6 Talar fractures and dislocations 45.3.7 Tarsal coalitions 45.3.8 Accessory navicular 45.4 Presurgical assessment 45.4.1 Clinical exam

701 702 705 705 705 706 706 706 706 706 707 707 707

Contents

45.5 Imaging 45.6 Goals in treatment 45.6.1 Treatment goals in flatfoot/cavus syndromes 45.6.2 Treatment goals in rheumatoid arthritis 45.6.3 Treatment goals in osteoarthritis 45.6.4 Treatment goals in calcaneal fractures 45.6.5 Treatment goals in talar fractures and dislocations 45.7 Corrective options 45.7.1 Distraction subtalar fusion 45.7.2 Lateral column lengthening 45.7.3 Double hindfoot fusion 45.7.4 Subtalar and talonavicular fusion 45.7.5 Triple arthrodesis 45.7.6 Pantalar arthrodesis 45.7.7 General complications 45.7.8 Postoperative management requirements in hindfoot fusions 45.8 Areas of future interest References

46. Biomechanics of Foot and Ankle Fixation

708 710 710 711 712 712 712 712 712 713 714 715 716 717 718 718 718 718

47.3.2 Bony and ligamentous ankle anatomy with arthroplasty 737 47.3.3 Ankle motion in the three cardinal planes with arthroplasty 737 47.3.4 Bony morphology and alignment with arthrodesis 738 47.4 Biomechanical considerations/ complications of arthroplasty or arthrodesis 738 47.4.1 Ankle alignment/malalignment 738 47.4.2 Gait mechanics 738 47.4.3 Arthritis at distal joints 738 47.4.4 Component wear or failure 739 47.4.5 Cadaveric gait simulation of arthroplasty and arthrodesis 739 47.4.6 Computational models of arthroplasty and arthrodesis 740 47.5 Biomechanical outcomes 741 47.6 Clinical outcomes 741 47.6.1 Safety 741 47.6.2 Effectiveness 742 47.6.3 Costs 743 47.6.4 Patient subgroups 743 47.7 Areas of future biomechanical research 744 References 744

721 48. Prosthetic Feet

Justin K. Greisberg and Christina E. Freibott 46.1 Introduction 46.2 Screws 46.3 Plates 46.4 Post and screw constructs 46.5 Nails 46.6 Beams 46.7 Areas of future research References

721 722 724 726 726 727 728 728

47. Ankle Arthroplasty and Ankle Arthrodesis

731

Daniel C. Norvell, Sagar S. Chawla and William R. Ledoux 47.1 Introduction 47.2 Brief description and history of surgical techniques 47.2.1 Ankle arthrodesis 47.2.2 Ankle arthroplasty 47.3 Biomechanical factors in presurgical assessment and consideration of arthroplasty or arthrodesis 47.3.1 Limb alignment with arthroplasty

xix

731 732 732 733

Glenn K. Klute 48.1 Introduction 48.2 Prescription and expected use of prosthetic feet 48.2.1 Activity levels 48.2.2 Activity bouts and durations 48.2.3 Activity in different environments 48.3 Form of prosthetic feet 48.3.1 Solid ankle cushioned heel 48.3.2 Fixed-angle stiffness 48.3.3 Variable-angle stiffness 48.3.4 Variable stiffness 48.3.5 Powered prosthetic feet 48.4 Function of prosthetic feet 48.4.1 Clinical trials 48.4.2 Mechanical property tests 48.4.3 Musculoskeletal modeling and simulation 48.5 Future prosthetic foot research References Index

736 736

749 749 750 751 751 751 752 753 753 754 755 755 755 756 758 759 759 760 765

Preface As researchers, it is of course in our best interests to emphasize what we do not know about the foot and ankle. Indeed, any future grant funding depends largely on our ability to convince reviewers that there are pressing questions that need to be answered about the complex assortment of bones and sinew attached to the end of the leg. Even though it is true that this anatomy has received far less attention than the knee or hip, for example, this may sometimes come at the expense of downplaying the progress that has been made in our understanding of the function of the foot and ankle; understanding that has come about through the dedication and hard work of countless individuals across the globe. Our hope is that this book will serve as a celebration of what we have learned about the foot and ankle over the past decades, whether the techniques that have been developed to study it, our understanding of diseases affecting it, or the interventions that can be performed to help with these problems. We would like to acknowledge and thank Joseph Iaquinto, PhD, for his work during the planning and initial editing of the book. Carrie Bolger provided an infinite well of patience and a steady, firm presence as we completed this project, as did many others at Elsevier. This book would of course not have been possible without the outstanding contributions from the chapter authors, and we owe them all a debt of gratitude. This debt may be discharged via coffee or beer at a conference some time in the future. William (Bil) would like to thank his wife (Kristin) and their three kids (Quinn, Kiernan, and Greysan) for always being there, as well as his dad (Bill) and father-in-law (Vaughn) for their continued interest and support. He would also like to acknowledge Bruce Sangeorzan, MD, and our colleagues at CLiMB, many of whom not only contributed the content for the book but also fostered the community of scientific research that made this book possible. Scott would like to thank his family, especially Michelle and Talia, for their support (“Wait, you’re still working on that?”) and understanding. He would also like to acknowledge all past and present collaborators for keeping the job interesting.

William R. Ledoux Scott Telfer

xxv

Anatomical Terms Used in Foot and Ankle Biomechanics As we set out to edit this book, we realized that terminology was going to be an important concept that we had to address early on in the process. Historically, within our own laboratories and within the classroom, we have circulated lists of anatomical do’s and don’ts related to the foot and ankle. Those lists formed the starting point for a document that we circulated to prospective chapter contributors, which after much feedback from Tom Greiner (Chapter 3) and Bruce Hirsch (Chapter 1), led to the following detailed descriptions. (See Chapter 3 for an in depth discussion of some of the more nuanced points.) While the list is not meant to address the full spectrum of potentially overlapping or confusing terms associated with foot and ankle biomechanics (and there are many), it is intended to provide a level of consistency among the chapters in this book, and we have made every endeavor to ensure that the chapters were edited accordingly. After organizing as suggested at the end of Chapter 3, we now present this list for the reader to better understand the multitude of concepts described in this book. G

G

G

G

G

The use of eponyms. G In general, avoid eponyms (Lisfranc joint, Chopart joint, etc.). G Except for the Achilles tendon, which is permissible. G For clarity, we have sometimes allowed eponyms parenthetically. Names for limb segments and other relevant anatomy. G Use “leg” as opposed to “lower leg” or “shank” for part of the lower extremity between the ankle and knee. G “Hindfoot” refers to the calcaneus and talus; “midfoot” refers to the navicular, cuboid, and cuneiforms; “forefoot” refers to the metatarsals and toes. G Use “hindfoot” not “rearfoot.” G Use “neutrally aligned” not “rectus.” G Avoid using “pes planus foot type,” as that literally means “foot flat foot type.” Instead, use “pes planus” or “planus foot type.” G “Fibularis” is an acceptable alternative to “peroneal.” (Chapter 3 uses the former, and the rest uses the latter.) Description of how the anatomical position is defined for the foot and for the foot and leg, and how those definitions relate to the body as a whole. G In general, “plane” refers to “leg-based plane” as opposed to “foot-based plane” (e.g., ankle plantarflexion and dorsiflexion are defined in the “leg-based” sagittal plane or just the sagittal plane, Fig. 1, modified from Fig. 3.2). G In general, we discuss the foot with the ankle at 90 degrees with the foot midline orientated parallel to the body midline (i.e., neutral position). This will be referred to in the text as “foot perpendicular” or “neutrally orientated.” This is opposed to the ankle fully plantarflexed, or “foot parallel,” which will not be used in this book (outside of being reviewed in Chapter 3). Note this is not “comfortable stance,” which usually entails some amount of external foot rotation. Note also that if we are talking about the metatarsophalangeal joints or phalanges, the planes change from leg-based to foot-based. In this case, hindfoot varus or valgus is in the leg-based coronal plane while hallux varus or valgus is in the foot-based coronal plane. How the anatomical planes are defined for the foot, and how they are related to the body as a whole. G These relationships are defined as below (Fig. 2, modified from Fig. 3.5). However, there is inconsistency in the literature about the foot-based coronal and transverse planes when the ankle is perpendicular. To address this, we have explicitly described motions below. G Use coronal rather than frontal for the plane that divides the body into anterior and posterior; either is fine, but we are using coronal to be consistent. How anatomical directions need to be defined relative to the foot and to the body. G These relationships are demonstrated below (Fig. 3, modified from Figs. 3.1 and 3.4). xxvii

xxviii

Anatomical Terms Used in Foot and Ankle Biomechanics

FIGURE 1 Foot and leg planes.

FIGURE 2 Body and foot planes.

G

How movements will be defined according to the definitions provided earlier and if some movement terms need to be considered joint specific or apply to the foot as a whole. G When reporting intersegmental kinematics, as far as possible, movements should be described with reference to the proximal segment (i.e., movement of the whole foot should be described in the reference frame of the leg). There are several approaches used to define these reference frames, and these should be clearly stated or the primary source cited. G Avoid the use of pronation/supination; instead discuss motion/position in specific cardinal planes. There are many issues with the use of pronation/supination, but two key ones are the following: (1) a “pronated” foot has a forefoot that is supinated relative to hindfoot and (2) pronation/supination is sometimes used in literature to describe just coronal plane motion. (See Chapter 3 for a more detailed consideration.) G Sagittal plane motion of the entire foot relative to the leg, that is, at the ankle joint, is referred to as plantarflexion/dorsiflexion (Fig. 4). Sagittal plane motion of other individual foot joints is referred to as flexion or

Anatomical Terms Used in Foot and Ankle Biomechanics

xxix

FIGURE 3 Anatomical directions.

FIGURE 4 Motion of foot relative to leg. Sagittal medial view of a left foot and leg (left). Ankle posed in neutral flexion (white solid outline), plantarflexion (red dotted outline), and dorsiflexion (green dotted outline). Coronal posterior view (center) of a neutral left foot (white solid outline) with an inverted heel (red dotted outline) and an everted heel (green dotted outline). Transverse superior view (right) of a neutral left foot (white solid outline) with an internally rotated foot (red dotted outline) and externally rotated foot (green dotted outline).

xxx

Anatomical Terms Used in Foot and Ankle Biomechanics

G

extension. Coronal plane motion of the entire foot is referred to as inversion/eversion. Coronal plane motion of other individual foot joints is defined below, as are select coronal plane joint poses. Transverse plane motion of the entire foot is referred to as internal/external rotation. Transverse plane motion of individual foot joints is defined below. Note: “plantarflexion” and “dorsiflexion” are both one word.

FIGURE 5 Hindfoot motions. Sagittal medial view of neutral ankle position (left top) with a plantarflexed talus and calcaneus (left bottom, red) and a dorsiflexed talus and calcaneus (left bottom, green). Coronal posterior view of neutral ankle position (center top) with an inverted calcaneus (center bottom, red) and an everted calcaneus (center bottom, green). Transverse superior view of neutral ankle position (right top) with an internally rotated talus and calcaneus (right bottom, red) and an externally rotated talus and calcaneus (right bottom, green).

Anatomical Terms Used in Foot and Ankle Biomechanics

G

G

xxxi

Hindfoot (i.e., calcaneus to tibia, calcaneus to talus, talus to tibia, Fig. 5) sagittal plane motion 5 plantarflexion/dorsiflexion coronal plane motion 5 inversion/eversion coronal plane joint pose 5 varus/valgus transverse plane motion 5 internal/external rotation Midfoot to hindfoot (e.g., navicular to talus, cuboid to calcaneus, Fig. 6) Note: these are described in terms of the leg-based coordinate systems. sagittal plane motion 5 flexion/extension coronal plane motion 5 inversion/eversion coronal plane joint pose 5 varus/valgus

FIGURE 6 Midfoot to hindfoot motions. Sagittal medial view of a neutral navicular and cuboid position (left top) with a flexed navicular and cuboid (left bottom, red) and an extended navicular and cuboid (left bottom, green). Coronal posterior view of a neutral navicular and cuboid position (center top) with an inverted navicular and cuboid (center bottom, red) and an everted navicular and cuboid (center bottom, green). Transverse superior view of a neutral navicular and cuboid position (right top) with an internally rotated navicular and cuboid (right bottom, red) and an externally rotated navicular and cuboid (right bottom, green).

xxxii

G

Anatomical Terms Used in Foot and Ankle Biomechanics

transverse plane motion 5 internal/external rotation Forefoot to midfoot (e.g., first metatarsal to medial cuneiform, fifth metatarsal to cuboid, Fig. 7) Note: these are described in terms of the leg-based coordinate systems. sagittal plane motion 5 flexion/extension coronal plane motion 5 inversion/eversion coronal plane joint pose 5 varus/valgus transverse plane motion 5 internal/external rotation

FIGURE 7 Forefoot to midfoot motions. Sagittal medial view of a neutral first metatarsal position (left top) with a flexed first metatarsal (left bottom, red) and an extended first metatarsal (left bottom, green). Coronal posterior view of a neutral first metatarsal position (center top) with an inverted first metatarsal (center bottom, red) and an everted first metatarsal (center bottom, green). Transverse superior view of a neutral first metatarsal position (right top) with an internally rotated first metatarsal (right bottom, red) and an externally rotated first metatarsal (right bottom, green).

Anatomical Terms Used in Foot and Ankle Biomechanics

G

xxxiii

Forefoot to hindfoot (e.g., first metatarsal to talus, Fig. 8) Note: these are described in terms of the leg-based coordinate systems. sagittal plane motion 5 flexion/extension coronal plane motion 5 inversion/eversion coronal plane joint pose 5 varus/valgus transverse plane motion 5 internal/external rotation

FIGURE 8 Forefoot to hindfoot motions. Sagittal medial view of a neutral first metatarsal position (left top) with a flexed first metatarsal (left bottom, red) and an extended first metatarsal (left bottom, green). Coronal posterior view of a neutral first metatarsal position (center top) with an inverted first metatarsal (center bottom, red) and an everted first metatarsal (center bottom, green). Transverse superior view of a neutral first metatarsal position (right top) with an internally rotated first metatarsal (right bottom, red) and an externally rotated first metatarsal (right bottom, green). Other metatarsals follow the same motion as the first.

xxxiv

G

Anatomical Terms Used in Foot and Ankle Biomechanics

Phalanges to metatarsals (e.g., proximal first phalange to first metatarsal—switch to foot-based planes) (Fig. 9). Note: these are described in terms of the foot-based coordinate systems. sagittal plane motion 5 flexion/extension coronal plane motion 5 adduction/abduction coronal plane joint pose 5 varus/valgus transverse plane motion 5 inversion/eversion

FIGURE 9 Phalange to metatarsal motions. Sagittal medial view of a neutral first proximal phalanx position (left top) with a flexed first proximal phalanx (left bottom, red) and an extended first proximal phalanx (left bottom, green). Coronal (leg-based) or transverse (foot-based) posterior view of a neutral first proximal phalanx position (center top) with an inverted first proximal phalanx (center bottom, red) and an everted first proximal phalanx (center bottom, green). Transverse (leg-based) or coronal (foot-based) superior view of a neutral first proximal phalanx position (right top) with adducted first proximal phalanx (right bottom, red) and abducted first proximal phalanx (right bottom, green).

Anatomical Terms Used in Foot and Ankle Biomechanics

xxxv

William R. Ledoux RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States Department of Mechanical Engineering, University of Washington, Seattle, WA, United States Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Scott Telfer Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States Department of Mechanical Engineering, University of Washington, Seattle, WA, United States RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States

Joseph M. Iaquinto RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States Department of Mechanical Engineering, University of Washington, Seattle, WA, United States Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Chapter 1

Anatomy of the Foot Bruce Elliot Hirsch* Department of Neurobiology and Anatomy, Drexel University College of Medicine, Philadelphia, PA, United States

Abstract The anatomy of the foot and ankle is quite complex, and entire books have been written on the topic. For the purposes of this chapter, an abridged anatomical review is provided. The structures described are limited to those within or in contact with the foot, except for the extrinsic muscles, whose bellies and tendons are located in the leg. This includes descriptions of the skeletal structures of the foot, such as the bones and joints, and covers ligaments as well. Next, the muscles and tendons of the foot are described in some detail, before the nerves and blood vessels are reviewed.

When discussing the functional anatomy of the foot, it is important to note that what happens there is strongly influenced by and dependent on many proximal structures. It is not possible to include much information on such structures in this chapter; however, it is impossible to describe and understand its anatomy without considering the relationships and interactions with certain more proximally located structures.

1.1

Skeletal structures

The regions of the tibia and fibula, which are involved in forming the ankle and inferior tibiofibular joint, are described here, in addition to the actual foot bones. Certain more proximal areas of the leg bones and the femur, where some of the extrinsic foot muscles arise, will be mentioned when those muscles are described.

1.1.1 Tibia and fibula The distal end of the tibial shaft expands to rest on the superior surface of the talus. Its medial side projects distally along the side of the talus, forming the medial malleolus, which is a wall of the mortise holding the talus in place. The anterior and posterior edges of the malleolus meet at a point called the anterior colliculus. The lateral side of the shaft is grooved longitudinally, to help hold the fibula in place. Inferiorly, the fibular shaft is secured against the lateral tibial surface by an interosseous ligament to complete the mortise. The narrow fibular shaft expands at its distal end to form the upside-down, pyramidal, lateral malleolus, whose apex extends 1 to 2 cm more distally than the medial malleolus. Its posterior border is grooved for the fibular muscle tendons. These two malleoli create the mortise which receives the talar body, where it is free to dorsiflex and plantarflex.

1.1.2 Segments of the foot skeleton The skeleton of the foot (Figs. 1.1, 1.2 and 1.3) includes seven anatomically short bones, making up the tarsus: the talus, calcaneus, navicular, cuboid, medial cuneiform, intermediate cuneiform, and lateral cuneiform. The first two are referred to as the hindfoot bones and the rest form the midfoot. They are chunky, each having one constant centrally located ossification center, and possibly others. They have limited motion individually but often work in conjunction with other bones to allow complicated directional control. *Retired. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00040-8 © 2023 Elsevier Inc. All rights reserved.

3

4

PART | 1 Introduction

Distal phalanx Middle phalanx

Proximal phalanx II I

III Head of phalanx

IV

Shaft of phalanx V

Head of phalanx

Base of phalanx

Shaft of phalanx

Base of phalanx Tarsometatarsal joints (LISFRANC's joint)

Lateral cuneiform Medial cuneiform Tuberosity of fifth metatarsal bone

Intermediate cuneiform

Cuboid Transverse tarsal joint (CHOPART's joint)

Navicular Head

Talus

Lateral process of talus

Trochlea of talus

Calcaneus FIGURE 1.1 Foot skeleton, dorsal view. Used with permission Paulsen F, Waschke J. Sobotta atlas of human anatomy. In: General anatomy and musculoskeletal system, vol.1, 15th ed. ISBN: 9780702052514.

The forefoot is made up of the metatarsals and phalanges, eighteen or nineteen anatomically long bones, even though some of them are short enough to be mere nubbins. Each comprises an elongated shaft, which ossifies from a primary ossification center, enlarged ends (at least one of which is articular); and a constant secondary ossification

Anatomy of the Foot Chapter | 1

5

Tuberosity of distal phalanx

Distal phalanx Middle phalanx I

Proximal phalanx

II III Sesamoid bones IV

Phalanges V

Tuberosity of first metatarsal bone

Metatarsals I–V

Medial cuneiform Intermediate cuneiform Tuberosity of fifth metatarsal bone

Lateral cuneiform Groove for fibularis longus tendon Tuberosity

Tuberosity

Navicular

Cuboid Head of talus

Calcaneus

Talus

Lateral process of calcaneal tuberosity

Sustentaculum tali

Medial process of calcaneal tuberosity

FIGURE 1.2 Foot skeleton, plantar view. Used with permission Paulsen F, Waschke J. Sobotta atlas of human anatomy. In: General anatomy and musculoskeletal system, vol.1, 15th ed. ISBN: 9780702052514.

center at one end (and possibly another). The five metatarsals lie side by side, articulating with all the midfoot bones except the navicular. At their distal ends the metatarsals articulate with the proximal phalanges, and they in turn (except in the hallux) articulate with the middle phalanges. The hallucal proximal phalanx and the middle phalanges articulate with the distal phalanges.

6

PART | 1 Introduction

Tarsal bones

Transverse tarsal joint (CHOPART's joint) Navicular Intermediate cuneiform Trochlea of talus Neck

Talus

Lateral cuneiform Tarsometatarsal joints (LISFRANC's joint) Metatarsals

Lateral malleolar facet Posterior process of talus

Phalanges

Tarsal sinus Calcaneus Fibular trochlea Calcaneal tuberosity

Cuboid

Tuberosity of metatarsal V

FIGURE 1.3 Foot skeleton, lateral view. Used with permission Paulsen F, Waschke J. Sobotta atlas of human anatomy. In: General anatomy and musculoskeletal system, vol.1, 15th ed. ISBN: 9780702052514.

1.1.3 Talus The talus is the second largest bone in the foot, after the calcaneus (Fig. 1.1). It articulates with the tibia superiorly, the tibia and fibula at its sides, and the calcaneus and navicular inferiorly and anteriorly. The superior surface of the body is entirely articular, pulley-like, and trapezoidal in outline, and usually 3 5 mm wider anteriorly than posteriorly. At its sides it is continuous with the articular areas of the medial and lateral surfaces. The lateral malleolar surface of the body contains a large inverted triangular facet for the lateral malleolus. The medial malleolar surface of the body presents a comma-shaped facet at its superior border, wide anteriorly and tapering posteriorly. These three articular areas form a continuous surface which fits into the ankle mortise. The inferior surface of the bone contains a groove, the sulcus tali, which faces a similar groove on the calcaneus and houses the interosseous talocalcaneal ligament. The posterior surface of the body is small, consisting of two tubercles which define a groove that controls the path of the flexor hallucis longus tendon. The talar neck is a constricted area between the body and the head, making up about 30% of the bone’s length. The head is convex in all directions, with three articular areas separated by low ridges. The long axis of the largest facet, the anterior, is set at an angle of about 45 degrees to the horizontal, higher laterally. It articulates with the navicular. The two smaller facets articulate with the sustentaculum tali of the calcaneus.

1.1.4 Calcaneus The calcaneus is the most massive and the most posterior bone of the foot (Fig. 1.2). It lies inferior to the talus, with which it articulates, and is separated from the ground at its posterior end by skin and fat. It rests at an angle to the ground so that the anterior end is elevated. On the lateral side the anterior half or so of the bone appears scooped out, creating a hollow where the talus sits. Anteriorly, it articulates with the cuboid and occasionally the navicular. The superior surface can be divided into four areas: two rough and two articular. The large articular facet near the center of the surface is known as the posterior talar articular surface. The shelflike sustentaculum tali at the anteromedial corner of the bone can have either one long articular facet or two smaller facets adjacent to each other; in either case they are called the anterior and middle talar articular surfaces. The sulcus calcanei separates the two articular regions. The oblong plantar surface has a large bulge at its posterior end, the calcaneal tuberosity, which is divided into medial and lateral processes by a pronounced notch. The long plantar ligament runs from the tuberosity to almost the anterior end of the bone. The smooth anterior tubercle, an attachment site of the plantar calcaneonavicular ligament, is at the distal end of the surface.

Anatomy of the Foot Chapter | 1

7

The shape of the lateral surface has been described as reminiscent of an old shoe, larger posteriorly and with a strangely truncated toe box. It is generally rough, except for an anteroinferiorly directed groove lying inferior to the sulcus calcanei, which marks the path of the peroneus longus tendon. The medial surface is quadrilateral and generally smooth. It bears a thick shelflike projection of bone near its anterosuperior angle, the sustentaculum tali, which supports two articular facets of the talus. The inferior half or so of the ovoid posterior surface is rough, for the insertion of the calcaneal tendon. The upper part is smooth. The small anterior surface is roughly like an inverted triangle in outline, and wholly articular.

1.1.5 Navicular The navicular, roughly oval and lying horizontally, is located between the talus and the three cuneiform bones (Fig. 1.1). In addition, it usually articulates with the cuboid laterally and sometimes with the calcaneus. The posterior surface is ovoid, broader laterally than medially, concave, and wholly articular. The anterior surface is wholly articular and generally convex. Two slight, roughly vertical ridges divide it into three facets for the cuneiform bones. The dorsal surface is convex anteriorly. The plantar surface, facing inferiorly and slightly laterally, is quadrilateral. It joins the dorsal surface at its medial end, where they form the tuberosity. A groove separates the tuberosity from the rest of this surface. The lateral surface usually presents a small facet inferiorly for articulation with the cuboid.

1.1.6 Cuboid The cuboid is actually more pyramidal than cubic in its shape (Fig. 1.3). The four-sided medial surface forms the base of the pyramid and the lateral surface, which is very small, the apex. The medial surface has an oval articular facet for the lateral cuneiform near its center, and there is usually a second articular facet for the navicular. The lateral surface has a deep notch continuous with the fibular sulcus. The dorsal surface is trapezoidal, faces superolaterally, and is continuous with the lateral, not the dorsal, surface of the calcaneus. The plantar surface is the largest surface of the cuboid and has an outline similar to that of the dorsal. A deep groove, the sulcus for the tendon of peroneus longus, runs parallel to the anterior border of this surface. The sulcus is bounded posteriorly by a prominent rounded ridge of bone, the promontory, whose bulbous lateral end is known as the tuberosity of the cuboid. The tuberosity usually bears a small smooth facet for a sesamoid in the peroneus longus tendon. The anterior surface is roughly triangular in outline, with its apex facing laterally; the lateral angle, however, is frequently quite rounded. The surface is wholly articular, divided by a low vertical ridge into two adjacent facets for the fourth and fifth metatarsal bones. The posterior surface is somewhat kidney shaped and wholly articular with the anterior calcaneal surface.

1.1.7 Medial cuneiform The medial is the largest of the three cuneiform bones (Fig. 1.1). It is a short bone, wedge-shaped with its base plantar. The four-sided medial and lateral surfaces share a border, which forms the dorsal crest. The medial and lateral surfaces are higher at their anterior ends than at the posterior, so the crest is arched. The medial surface has a shallow groove running from its posterosuperior angle to the anteroinferior corner, where there is a small smooth area. The tibialis anterior tendon runs in this groove to an insertion on that area. On the lateral surface a facet shaped like an inverted L is situated along the superior and posterior borders of the bone, for the middle cuneiform. The anterior surface is reniform, with its hilum located laterally. This surface is convex and wholly articular. The posterior surface, smaller than the anterior, is more or less pear-shaped with the stem end pointing upwards. It is concave and wholly articular. The plantar surface is quadrilateral and rough. A bulge, the tuberosity of the medial cuneiform, is present near its posterior end.

1.1.8 Intermediate cuneiform The intermediate cuneiform is the smallest of the three cuneiform bones (Fig. 1.1). It is wedge-shaped, with its base forming the dorsal surface and the crest facing plantarly. It articulates with the medial cuneiform, the lateral cuneiform,

8

PART | 1 Introduction

the navicular, and the second metatarsal. Its posterior end is in line with the other two cuneiforms, but the anterior end doesn’t extend as far forward as they do. This creates a recess into which the second metatarsal base fits. The anterior and posterior surfaces are each about the same size, triangular, and wholly articular. The anterior surface is slightly convex, while the posterior surface is flat or slightly concave. The medial and lateral surfaces are both four sided. The medial presents an inverted L-shaped facet along its posterior and superior borders. The lateral surface bears an elongated vertical facet along its posterior border. This facet is wider superiorly and usually has a waist near its center. The roughened dorsal surface is almost square, but slightly narrower anteriorly than posteriorly. The crest, the inferior border, is a rounded ridge, slightly thicker at its posterior end.

1.1.9 Lateral cuneiform The lateral cuneiform, intermediate in size between the two other cuneiforms, is wedge-shaped with its base up (Fig. 1.1). It has five named surfaces and a crest. It articulates with the intermediate cuneiform, the cuboid, navicular, and second and third metatarsals, and sometimes with the fourth metatarsal base. Both the anterior and posterior surfaces are triangular, base up, and apex down. The anterior surface is wholly articular and slightly concave. The posterior surface bears a flat or slightly concave articular facet, a blunted triangle in shape, in its superior part, and a rough area inferiorly. The facet is continuous with the articular facets on the sides of the bone. The medial and lateral surfaces are each four sided. The medial bears an elongated facet for articulation with the intermediate cuneiform, and sometimes one or two facets for the second metatarsal. The lateral surface has two articular facets, a large one at its posterosuperior angle and a smaller one at the anterosuperior corner, for the cuboid and fourth metatarsal respectively. The rectangular dorsal surface faces almost directly superiorly. The medial and lateral surfaces meet to form the crest, the inferior border.

1.1.10 Metatarsals The metatarsus comprises five small anatomically long bones numbered 1 through 5, beginning with the medial one (Fig. 1.1). They are in contact at their bases but spread out slightly anteriorly, so the width of the metatarsus where it joins the phalanges is greater than where it articulates with the tarsus. The proximal end of each bone is superior to the distal, so that only the heads, the distal ends, contact the substrate. The metatarsals all share certain features. Each bone has an expanded proximal end, the base; a hollow shaft; and an expanded distal end, the head. The bases are more or less pyramidal with flattened posterior surfaces by which each articulates with one of the tarsal bones. The sides of the bases are flattened where they adjoin neighboring bones. Each shaft has three surfaces and three borders, usually named by their orientation at their proximal ends because they may twist as they extend distally. One surface of each shaft faces more or less dorsally and the other surfaces face sideways or plantarly. Shortly posterior to the articular area covering the head of the bone there are two tubercles, one on each side, dorsally. Sometimes reference is made to the anatomical neck, which is a narrow region where the head and shaft meet, proximal to the tubercles, or to the surgical neck, at the groove between the tubercles and the head.

1.1.10.1 First metatarsal The first metatarsal is the shortest, thickest, strongest, and most massive of all the metatarsal bones. It articulates with the medial cuneiform proximally, and with the proximal phalanx of the first ray and the sesamoids of flexor hallucis brevis distally. Although it is in contact with the second metatarsal base there is no true joint between them. The base has posterior, dorsal, lateral, and medial surfaces. The posterior surface, which articulates with the medial cuneiform, presents a vertical reniform articular facet whose hilum is directed laterally. This surface is wholly articular. The medial surface, which is in line with the medial surface of the medial cuneiform, bears a small tubercle near its center marking the insertion of tibialis anterior. The first is the only metatarsal with no articular facets on its sides, although it has pressure facets which can look very much like them. The medial and lateral surfaces of the base meet plantarly, forming a tuberosity where peroneus longus inserts. The shaft is triangular in cross section. The lateral surface lies almost vertically, and is very slightly concave in its short axis. The dorsal surface is continuous with the medial surface of the base, but it twists to face dorsally as it

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reaches the head. The plantar surface faces inferomedially at its proximal end and rotates to face inferiorly at the metatarsal head. The head of the first metatarsal, far larger than the others, is rounded, with the radius of the side-to-side curvature greater than that of the vertical. Its transverse shape changes from dorsal to plantar: it is strongly curved at its dorsal end, somewhat flatter anteriorly; and at its plantar end it is marked with two grooves separated by a ridge. Sesamoid bones of flexor hallucis brevis slide in these grooves.

1.1.10.2 Second metatarsal The longest metatarsal, projecting farthest anteriorly, is the second. Its base is inset between the medial and lateral cuneiforms proximally, articulating with them, the intermediate cuneiform, and the third metatarsal. Although it touches the first metatarsal, possibly separated from it by a bursa, there is no true joint. The base is pyramidal, presenting dorsal, medial, lateral, and posterior surfaces. The wholly articular posterior surface is triangular in outline, base up and apex down. The medial surface presents a facet at its posterosuperior angle to articulate with the medial cuneiform. In most bones this facet is continuous with the large facet on the posterior surface. Anteroinferior to this there is a smooth area which is a pressure facet where the first metatarsal rubs. The lateral surface presents two articular facets, superior and inferior. Each is divided by a small vertical ridge into two smaller articular areas for the articulations with the lateral cuneiform and third metatarsal. The lateral and medial surfaces of the base meet at the rough, rounded inferior border. The shaft tapers from the base toward the head, presenting three surfaces. It has an apparent torsion as though the distal end was slightly everted, but the surfaces are named by their positions at their proximal end. The second metatarsal head is taller than its width. It presents a curved articular surface which covers the distal end of the bone, extending a short distance dorsally and a longer distance plantarly. It terminates plantarly in two posteriorly directed projections separated by a notch.

1.1.10.3 Third metatarsal The third metatarsal is somewhat shorter than the second and a little longer than the fourth. Its base is similar in shape to that of the second metatarsal, although the articular facets are arranged differently on the sides. The medial surface presents two facets along its posterior edge, the dorsal almost always the larger. The lateral surface bears a single articular facet at its posterosuperior angle. Where the medial and lateral surfaces of the base meet inferiorly, they form the wide, roughened inferior border. Its posterior surface articulates with the anterior surface of the lateral cuneiform. The shaft of the third metatarsal is very similar to that of the second except that it is shorter, and it is more distinctly bowed along its long axis so that the concavity faces inferolaterally. The head of the third metatarsal bone is also very similar to that of the second.

1.1.10.4 Fourth metatarsal The fourth metatarsal bone is slightly shorter than the third, and like the third it is slightly bowed along its long axis. However, the concavity faces more laterally. Like the second and third, this shaft also has a torsion to it so that the distal end seems slightly everted. The base of the fourth metatarsal has dorsal, medial, lateral, posterior, and plantar surfaces, and is close to cuboidal. The posterior surface of the base is vertically rectangular, almost plane, and wholly articular for the medial end of the anterior cuboidal surface. The medial surface presents a facet near its superior border for the third metatarsal. In those feet where the lateral cuneiform projects farther anteriorly than the cuboid there may also be another, smaller, articular facet for that bone. A rather large triangular facet for the fifth metatarsal base is located at the posterosuperior angle of the lateral surface. The shaft is similar to those of the second and third bones, except that it is shorter and stockier. The head is also similar to the heads of those two bones.

1.1.10.5 Fifth metatarsal The fifth metatarsal is shorter than the fourth. Its base is pyramidal, wider than its height, with dorsal, plantar, medial, and posterior surfaces. The posterior surface of the base is an elongated triangle with its apex on the lateral side, and wholly articular. The medial surface bears a rather large triangular facet which matches a facet on the lateral side of the

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fourth metatarsal. The dorsal and plantar surfaces are both rough; the plantar faces inferiorly and the dorsal superolaterally. They unite in a bulbous projection which extends posterolaterally, known as the tuberosity. The base leads smoothly into the shaft. The shaft presents three surfaces, medial, dorsal, and plantar, each of which is continuous with the similarly named surface of the base. It tends to be flattened dorsoplantarly. The fifth metatarsal head is very similar to those of the second, third, and fourth metatarsals, except that the articular surface extends even farther on the lateral side.

1.1.11 Phalanges Each toe has three phalanges except for the hallux, which has only two (Fig. 1.1). The phalanges form proximal, middle, and distal rows; and within each toe the sizes of the phalanges decrease from proximal to distal. In the lateral toes the phalanges within each row tend to decrease slightly in size from medial to lateral. The phalanges of the great toe are much larger than those of the other digits. Despite their small size the phalanges are long bones, each ossifying from a diaphysis and a single epiphysis located at the base. Every phalanx has a base, a shaft, and a head, although the smaller bones may be so reduced that they appear to be totally lacking a diaphysis. The positions of the phalanges are variable. Generally, the proximal phalanges lie horizontally or are very slightly tilted superiorly toward their heads. The heads of the middle and distal rows are pointed slightly inferiorly so that the tuberosities of the distal rows approach the ground.

1.1.11.1 Proximal phalanges The proximal phalanges of all the toes are similar in form. However, the proximal phalanx of the hallux is by far the most massive of the phalanges, reflecting its role in those parts of gait when much of the body is transmitted along the great toe. The shafts of the four lateral phalanges may all be of similar width or they may become thinner going laterally. The proximal phalanges are the longest bones in each toe. In the four lateral toes they are longer than the combined length of the middle and distal phalanges. Each has a base, a shaft, and a head. The posterior articular surface of each base is in the form of an oval with a flat inferior edge, or a very flattened triangle. It has only a shallow concavity. The plantar surface of the base is roughened and bears relatively large tubercles at the sides. The shafts are narrowest just distal to the center of the bone. They are flattened dorsoplantarly, although the lateral ones tend to be more cylindrical. The shafts thus have superior and inferior surfaces separated by medial and lateral borders. The superior surfaces are flat or very slightly convex along their long axes, but plantarly they are markedly arched. The superior and inferior surfaces are usually separated by wide medial and lateral borders. However, sometimes they may just be ridges where the skin ligaments attach. The pulley-shaped heads of the proximal phalanges are smaller than the bases. Each presents a trochlear articular surface with a distinct sagittal groove, which is visible as only a small crescentic facet dorsally and a larger area plantarly. Passing over the distal end of the bone, the articular area widens as it reaches the plantar surface. The median grooves in the articular area, which receive the crests in the middle phalangeal bases, also widen plantarly. The sides of the head are fairly flat, with small dorsal tubercles. In the hallux the anterior end of the bone is deviated laterally about 3 4 degrees.

1.1.11.2 Middle phalanges Only the lateral four toes have middle phalanges. They are much shorter than the proximal phalanges. The posterior surface of the base presents a large oval articular area whose long axis lies horizontally. A low vertical crest divides this area into two concave oval facets of equal size. This median crest is often elongated dorsally, extending far enough posterosuperiorly to rest on top of the head of its proximal phalanx. Inferiorly, tubercles are found on the sides of the base. The dorsal and plantar surfaces of the base are roughened. The shafts are considerably shorter and flatter than those of the proximal row, and these are the narrowest part of the bone. Each is thickest posteriorly where it joins the base, and, when viewed from the side, tapers anteriorly toward the head like a wedge. Each shaft thus has dorsal and plantar surfaces, and medial and lateral borders. The heads of the middle phalanges are similar to those of the proximal row, except that they do not extend as far plantarly, and their median grooves are relatively shallow.

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1.1.11.3 Distal phalanges The distal phalanges are the second phalanx of the hallux and the third phalanges of the four lateral toes. The distal phalanx of the great toe is a large, strong, broadened bone, capable of bearing much weight. Those of the other rays are very much reduced; the distal phalanx of the fifth toe is the smallest of all, and is sometimes fused with the middle phalanx. Except for their sizes the four lateral bones have a similar morphology. The distal phalanges are tiny long bones. Each is composed of a base, a shaft so short that it may be virtually nonexistent, and a distal extremity known as the tuberosity. The bases of these bones in the lateral four toes are very similar, except for size, with those of the middle phalanges. Their posterior surfaces, which are wholly articular, are oval with the long axis running transversely. A blunt median crest divides each surface into two connected, concave articular facets. The sides of each base may extend outward as points, or even as hooks which point anteriorly. The shafts of the distal phalanges are almost nonexistent, and may be described as the constricted area between the base and the tuberosity. Occasionally the base and the head are so close together that it is hard to actually see an area that looks like the shaft. The expanded distal ends are called the tuberosities. They are crescentic, oriented from side to side, flattened, and rougher on their plantar surfaces than they are dorsally. The roughness can extend onto whatever bit of shaft may be seen. The distal phalanx of the hallux is considerably larger than any of the others. Viewed from the side, it tapers to become much thinner near its tip. The wholly articular posterior surface of the base is a very elongated oval, with its long axis running transversely. The central median crest is a gentle elevation, not a sharp ridge. On either side of the crest the articular surface is concave and the tubercles on the sides of the base are large. The dorsal and plantar surfaces are rough. The distal hallucal phalanx is set at an angle to the axis of the metatarsal and proximal phalanx, veering laterally about 15 degrees. Part of this angulation may be at the interphalangeal joint, but most appears to be within the bone. The plantar surface of the hallucal shaft is rough, and often has a raised V-shaped ridge whose apex points distally and slightly laterally, where flexor hallucis longus inserts. The tuberosity is similar to those of the lesser toes, but much larger.

1.2

Joints

The joints described here are those which directly involve the bones of the foot.

1.2.1 Tibiofibular syndesmosis The inferior tibiofibular joint, the articulation between the distal ends of the tibia and fibula, is a fibrous joint. It lies between the fibular notch of the tibia and the triangular area within the inferior end of the interosseous fibular border. The interosseous ligament of the tibiofibular syndesmosis is formed of many short, strong bands which firmly tie the adjacent bones together. They form a short ligament of large cross-sectional area which prevents the forces at the ankle joint from wedging the talus upward between the leg bones. The anterior inferior tibiofibular ligament of the syndesmosis is a bundle of fascicles, most often three, running inferolaterally between the anterior border of the fibular notch of the tibia and the anterior border of the lateral malleolus (Fig. 1.4). The posterior inferior tibiofibular ligament also runs inferolaterally from the distal end of the tibia. It extends from the posterior surface to the medial edge of the lateral malleolar sulcus, and is stronger than the anterior. The deepest and most inferior part of the posterior inferior tibiofibular ligament is known as the inferior transverse ligament. It bridges the space between the inferior ends of the tibia and fibula, thus forming a posterior wall to the ankle mortise.

1.2.2 Ankle joint The anatomic ankle joint is the talocrural joint, at which the trochlea of the talus fits into a socket formed by the distal ends of the tibia and fibula, the anterior inferior tibiofibular ligament, and the inferior transverse ligament. As a ginglymus or hinge joint the ankle’s motion is mostly restricted to dorsiflexion and plantarflexion. However, small amounts of rotation, adduction, and abduction are possible. The fibrous articular capsule surrounds the joint and is attached to the bones just at the edges of the articular surfaces. It is very thin anteriorly and posteriorly. Strong ligamentous structures reinforce the sides of the ankle, probably mixtures of both capsular and extracapsular fibers, forming collateral ligaments that on the medial side is known as the deltoid ligament because of its triangular shape, and on the lateral side it is simply called the lateral ligament. The deltoid ligament is stronger than the lateral (Fig. 1.5). It fans out from the medial malleolus to reach the navicular, calcaneus, and talus. It is considered to have four component parts, the deepest of which is the anterior tibiotalar

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Anterior tibiofibular ligament Lateral talocalcaneal ligament

Anterior talofibular ligament

Talocalcaneal interosseous ligament Dorsal cuboideonavicular ligament

Lateral malleolus

Dorsal cuneonavicular ligaments Dorsal metatarsal ligaments

Calcaneofibular ligament Calcaneal tendon

Deep transverse metatarsal ligaments

Long plantar ligament

Calcaneonavicular ligament

Calcaneocuboid ligament

Dorsal tarsometatarsal ligaments Fibularis [peroneus] brevis, tendon

Bifurcate ligament

FIGURE 1.4 Dorsal ligaments of the ankle and tarsal joints, lateral view. Used with permission Paulsen F, Waschke J. Sobotta atlas of human anatomy. In: General anatomy and musculoskeletal system, vol.1, 15th ed. ISBN: 9780702052514.

Fibula Posterior tibiotalar part

Medial [deltoid] collateral ligament

Tibia

Tibiocalcaneal part Anterior tibiotalar part Tibionavicular part

Tibialis posterior, tendon Tibialis anterior, tendon Posterior tibiofibular ligament

Plantar calcaneonavicular ligament

Plantar tarsometatarsal ligaments

Plantar calcaneonavicular ligament

Long plantar ligament

FIGURE 1.5 Plantar ligaments of the ankle and tarsal joints, medial view. Used with permission Paulsen F, Waschke J. Sobotta atlas of human anatomy. In: General anatomy and musculoskeletal system, vol.1, 15th ed. ISBN: 9780702052514.

ligament. It runs from the anterior border and anterior colliculus of the medial malleolus, to the rough area below the thick part of the “comma” on the medial surface of the talar body, or to the medial side of the neck. The other three ligaments are the posterior tibiotalar, the tibionavicular, and tibiocalcaneal. The tibiocalcaneal ligament attaches to the sustentaculum tali and merges with the edge of the plantar calcaneonavicular ligament. This lateral part is sometimes called the tibiospring ligament because it helps support the plantar calcaneonavicular ligament, thereby helping to maintain the talar head in position and protecting the longitudinal arch. The tibiotalar ligaments cross only the ankle joint.

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However, the tibiocalcaneal ligament also crosses the subtalar joint, and the tibionavicular ligament crosses the talocalcaneonavicular joint as well. The ligaments on the lateral side of the ankle form three cord-like bands called, as a group, the lateral ligament. The anterior talofibular ligament is the shortest of the three. It runs almost horizontally between the anterior border of the lateral malleolus and the lateral side of the talar body, and sometimes onto the neck. The posterior talofibular ligament is the strongest of these ligaments. It runs almost horizontally from the lateral malleolar fossa to the posterior rough strip on the lateral surface of the talus, extending all the way onto the lateral tubercle. This ligament, which helps separate the ankle and subtalar joints, is covered with synovial membrane on its superior and inferior surfaces. The calcaneofibular ligament is extracapsular. Its proximal attachment is the summit and the distal end of the lateral malleolus, and from there it passes posteroinferiorly to the fibular trochlea of the calcaneus. This ligament crosses perpendicular to the subtalar joint line, therefore affecting that joint as well as the ankle. Fibers of this ligament and the anterior talofibular ligament intermingle at the fibula. Even though some of the ankle ligaments cross the subtalar and talocalcaneonavicular joints, the synovial spaces are entirely separate. Fat intervenes between the fibrous and synovial capsular membranes anteriorly and posteriorly, where the synovial membrane is lax to allow for the movements at the ankle. There may be places where the fibrous capsule is deficient posteriorly, especially near the medial malleolus, so that the synovial membrane is unprotected.

1.2.3 Subtalar joint The subtalar or talocalcaneal joint is the articulation between the posterior calcaneal articular surface of the talar body and the posterior talar articular surface of the calcaneus. It does not include the articulation between the anterior and medial facets on these bones, which are part of the talocalcaneonavicular joint (Fig. 1.6). Functionally, it is somewhat Dorsal tarsometatarsal ligaments Metatarsal II

Metatarsal IV

Metatarsal I

Metatarsal V

Dorsal intercuneiform ligament

Tuberosity of fifth metatarsal bone

Dorsal cuneonavicular ligament

Dorsal cunecuboid ligament Cuboid Calcaneonavicular ligament Calcaneocuboid ligament

Navicular

Bifurcate ligament

Talocalcaneonavicular joint Fibularis [peroneus] brevis, tendon

Plantar calcaneonavicular ligament

Anterior talar articular surface Middle talar articular surface

Subtalar joint

Talocalcaneal interosseous ligament Posterior talar articular surface

Calcaneus

FIGURE 1.6 Dorsal ligaments of the ankle and tarsal joints. Used with permission Paulsen F, Waschke J. Sobotta atlas of human anatomy. In: General anatomy and musculoskeletal system, vol.1, 15th ed. ISBN: 9780702052514.

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artificial to consider the two joints separately even though they are distinct, because when there is movement at one joint there must also be movement at the other. That being said, sometimes, especially in clinical usage, the whole talocalcaneal joint complex is called the subtalar joint. Here the subtalar joint will be considered independently. The posterior facets on the talus and calcaneus are both roughly rectangular, although the facet on the talus is slightly longer and that on the calcaneus is somewhat wider. Along its long axis the posterior calcaneal articular surface of the talus is concave, receiving the convex facet on the calcaneus. The fibrous articular capsule completely surrounds the joint, and is attached close to the edges of the posterior calcaneal articular surface of the talus. It is also attached close to the medial and anterior sides of the articular surface on the calcaneus, but is a slight distance from it laterally and posteriorly. It consists of short fibers between the two bones. Various thickenings of the capsule form ligaments, although they tend to be somewhat indistinct. The posterior talocalcaneal ligament is usually Y-shaped. Its stem attaches to the posterior rough area on the superior surface of the calcaneus some distance from the articular facet, and its branches attach to the two tubercles on the posterior surface of the talus. The lateral talocalcaneal ligament, sometimes extremely weak and hard to define, is situated deep to the calcaneofibular ligament and reinforces the lateral side of the subtalar joint capsule. Its fibers parallel those of the calcaneofibular ligament, connecting the lateral talar process and the lateral surface of the calcaneus. The medial talocalcaneal ligament connects the medial tubercle of the talus to the posterior end of the sustentaculum tali and to the body of the calcaneus nearby. There are two extracapsular ligaments contributing to the joint. The interosseous talocalcaneal ligament lies within the tarsal canal, between the subtalar and talocalcaneal joint capsules. It is a flat band, about 15 mm from side to side, connecting the sulcus calcanei to the sulcus tali. The strong cervical ligament, connecting the talus and the calcaneus, lies lateral to the interosseous talocalcaneal ligament. It runs anterosuperomedially from the calcaneus to the neck of the talus. The synovial membrane lines the fibrous capsule. It is attached to the bones along the edges of the articular surfaces.

1.2.4 Talocalcaneonavicular joint At the talocalcaneonavicular joint the head of the talus fits into a socket formed by the calcaneus, the navicular, and the plantar calcaneonavicular ligament (Fig. 1.6). It is condyloid, but because part of the socket is of soft tissue it may give slightly to allow somewhat more mobility than would ordinarily be expected. Low lines or ridges separate the articular area of the talar head into various regions corresponding to the components of the socket. The head thus bears at least three recognizable articular areas, for the navicular, the plantar calcaneonavicular ligament, and for the anterior and middle calcaneal articular facets. The socket is a relatively large concavity, with its anterior part facing almost posteriorly and the posteroinferior part facing anterosuperiorly. The posterior surface of the navicular forms the anterior wall of the socket. The posterior part of the socket is made up of the anterior and middle talar articular facets on the calcaneus. These facets, which may be continuous with each other, are concave along their long axes, forming an arc on which the talus sits. There is a triangular gap, with its apex pointing laterally, between the two bony components of the joint socket. This space is closed by the plantar calcaneonavicular ligament, which forms the floor and medial side of the socket. Fibrocartilage covers most of the superior surface of the plantar calcaneonavicular ligament; and although the ligament is often called the spring ligament, it does not contain elastic tissue. In addition to the ligaments of its capsule, the talocalcaneonavicular joint is reinforced by the tibionavicular and tibiocalcaneal parts of the deltoid ligament on its medial side. Also, a strong ligament passes from the wide lateral end of the sulcus calcanei to the cuboid and the navicular, on the superolateral aspect of the foot. The ligament, known as the bifurcate ligament, has two parts, one of which is a ligament of the talocalcaneonavicular joint. This portion of the bifurcate ligament is the calcaneonavicular ligament, serving to reinforce the lateral side of the joint; the other portion is the calcaneocuboid ligament. Together, they attach to the calcaneus between the anterior talar articular facet and the cervical ligament and then, like a Y or a V, separate. Synovial membrane lines the fibrous capsule, including those soft tissues of the socket which are not covered with fibrocartilage.

1.2.5 Calcaneocuboid joint The sellar calcaneocuboid joint is formed by the posterior surface of the cuboid and the anterior surface of the calcaneus. It, along with the talocalcaneonavicular joint, forms a complex known as the transverse tarsal (Chopart’s) joint. The articular facet on the anterior surface of the calcaneus is a close fit to that of the cuboid.

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Although it is a sellar joint, the articular surfaces do not have the expected saddle shape. A saddle shape is a segment of the central region of an hourglass-shaped surface of revolution. The posterior surface of the cuboid, however, more closely matches one end of such a surface. Nevertheless, it still permits movement around two essentially perpendicular axes. The part of the cuboid which represents the waist of the surface of revolution extends posteriorly as the calcaneal process, and fits into a recess in the plantar surface of the calcaneus. The fibrous articular capsule is slightly thickened superiorly to form the dorsal calcaneocuboid ligament, which runs from the lateral surface of the calcaneus to the dorsal surface of the cuboid. The plantar calcaneocuboid ligament, wide and short, is also a thickening of the capsule (Fig. 1.7). Lying superior and parallel to the long plantar ligament, it extends from the anterior tubercle of the calcaneus to the cuboid. Barely 2 cm long and about the same in width, it spreads out on either side of the long plantar ligament. Two other ligaments reinforce the calcaneocuboid joint. The Metatarsophalangeal joints

Sesamoid bone

Deep transverse metatarsal ligament

Fibularis [peroneus] longus, tendon

Plantar metatarsal ligaments

Plantar tarsometatarsal ligaments

Tibialis anterior, tendon

Plantar tarsal ligaments Fibularis [peroneus] brevis, tendon Cuboid Plantar cuboideonavicular ligament

Tibialis posterior, tendon

Plantar calcaneonavicular ligament Plantar calcaneocuboid ligament

Calcaneal tuberosity

FIGURE 1.7 Deep plantar ligaments of the ankle and tarsal joints. Used with permission Paulsen F, Waschke J. Sobotta atlas of human anatomy. In: General anatomy and musculoskeletal system, vol.1, 15th ed. ISBN: 9780702052514.

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Collateral ligaments

Plantar ligaments

Deep transverse metatarsal ligament

Plantar tarsometatarsal ligaments

Medial cuneiform

Plantar cuneonavicular ligaments Plantar cuboideonavicular ligament

Navicular

Plantar calcaneocuboid ligament Plantar calcaneonavicular ligament Long plantar ligament Sustentaculum tali

Calcaneal tuberosity

FIGURE 1.8 Plantar ligaments of the ankle and tarsal joints. Used with permission Paulsen F, Waschke J. Sobotta atlas of human anatomy. In: General anatomy and musculoskeletal system, vol.1, 15th ed. ISBN: 9780702052514.

long plantar ligament is the longest ligament in the foot and one of the strongest in the body (Fig. 1.8). At its posterior end it attaches to the plantar surface of the calcaneus, covering the whole space between the tuberosity and the anterior tubercle. Most of the fibers attach to the promontory of the cuboid, but some of the more superficial continue anteriorly to form a roof over the fibular sulcus. They form the fibular sheath, enclosing the peroneus longus tendon. The most

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superficial fibers then continue on, attaching to the third and fourth, and sometimes the second or fifth, metatarsal bases. The bifurcate ligament includes a calcaneocuboid component, which attaches to the medial surface of the cuboid near its junction with the dorsal surface.

1.2.6 Cuboideonavicular joint The cuboideonavicular joint is sometimes a plane synovial joint and sometimes a syndesmosis. The synovial version allows very little motion because a short interosseous ligament binds the two bones closely. If the joint is synovial, an articular facet is located on the lateral surface of the navicular, at its anterior edge. The facet on the cuboid is at the posterior part of the medial surface, adjacent to the facet for the lateral cuneiform. If the joint is syndesmotic no articular facets will be present, and both surfaces will be entirely rough. Strictly speaking, a fibrous capsular layer only exists in a synovial joint. The ligaments, however, are always present whether or not the joint is synovial. Three ligaments reinforce the cuboideonavicular joint. The dorsal cuboideonavicular ligament attaches to the dorsal surface of the navicular, at its lateral side (Fig. 1.4). It runs anterolaterally, becoming wider and overlapping the corner of the lateral cuneiform, to reach the posteromedial part of the dorsal surface of the cuboid. The plantar cuboideonavicular ligament connects the plantar surfaces of the two bones. The interosseous cuboideonavicular ligament, short and strong, runs between the bones’ adjacent surfaces. In the synovial version of the joint the articular areas lie adjacent to the cuneocuboid and lateral cuneonavicular joints, and its synovial space is continuous with theirs.

1.2.7 Cuneonavicular joints The posterior surfaces of the three cuneiform bones articulate with their corresponding facets on the anterior surface of the navicular. These synovial joints share a common capsule which is also continuous with the capsules of the intercuneiform, cuneocuboid, and cuboideonavicular joints. The cuneonavicular joints are plane joints, and all their opposing bone surfaces are wholly articular. Only a small amount of movement occurs at these joints. The anterior surface of the navicular is divided by vertical ridges into three facets, one for each of the cuneiforms. The posterior surfaces of the cuneiforms lie one right against the next, and the curvatures of the articular areas correspond to those of the navicular. Their positions reflect the shape of the transverse pedal arch. The fibrous articular capsule is given over almost entirely to ligaments, which form a continuous sheet on the two surfaces and medial side of the foot. A fibrous capsule is lacking on the lateral side, where the lateral cuneonavicular joint is continuous with the cuneocuboid and (when it is present) cuboideonavicular joints. The medial dorsal cuneonavicular ligament is a strong band extending from the tuberosity of the navicular to the crest and medial surface of the medial cuneiform (Fig. 1.4). The intermediate and lateral dorsal ligaments run from the dorsal surface of the navicular to the corresponding cuneiforms. The three plantar cuneonavicular ligaments extend from the plantar surface of the navicular to the tuberosity on the plantar surface of the medial cuneiform, and to the crests of the intermediate and lateral cuneiforms. Expansions of the tibialis posterior tendon merge with and reinforce these ligaments. The synovial cavity of the three cuneonavicular joints is continuous with those of the intercuneiform and cuneocuboid joints, the intermediate and lateral cuneometatarsal joints, and the intermetatarsal joints among the second, third and fourth metatarsals. Together, these form the great tarsal synovial cavity, which will be further described in the section on the tarsometatarsal joints.

1.2.8 Intercuneiform and cuneocuboid joints The cuneiforms and cuboid form the most distinct part of the transverse arch, the intermediate cuneiform representing the arch’s highest point. To maintain the stability of the arch, the joints are reinforced by strong plantar and interosseous ligaments. They are plane synovial joints but little motion occurs among the bones. The cartilage covered articular facets occupy only part of each apposing surface; the remainder of the surfaces are rough for attachment of the interosseous ligaments. The medial intercuneiform joint is between the posterior part of the lateral surface of the medial cuneiform and the medial surface of the intermediate cuneiform. The lateral intercuneiform joint is between the lateral surface of the intermediate cuneiform and the medial surface of the lateral cuneiform. The lateral surface of the lateral cuneiform and the medial surface of the cuboid form the cuneocuboid joint. The fibrous articular capsules of these joints blend with the

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capsule of the cuneonavicular joint. Large areas of the capsules are thickened to form ligaments, but between those ligaments the capsules are thin and sometimes deficient. The dorsal intercuneiform and cuneocuboid ligaments pass from dorsal surface to dorsal surface, and to the crest of the medial cuneiform (Fig. 1.6). These ligaments are all rather weak. The much stronger plantar intercuneiform and cuneocuboid ligaments pass from bone to bone, linking the plantar surface of the medial cuneiform, the crests of the intermediate and lateral cuneiforms, and the plantar surface of the cuboid. The interosseous intercuneiform and cuneocuboid ligaments are short and strong, connecting the adjoining rough surfaces of the adjacent bones, and merging with the plantar ligaments inferiorly. The intercuneiform and cuneocuboid joints contribute to the great tarsal synovial cavity, which is described in the section on the tarsometatarsal joints.

1.2.9 Tarsometatarsal joints The tarsometatarsal joints, together commonly known as Lisfranc’s joint, are formed by the articulations between the cuneiforms and the cuboid on one side, and the bases of the five metatarsals facing them. They lie along an oblique line running posterolaterally from the medial cuneiform first metatarsal joint. The differences in anterior projection among the cuneiforms and cuboid create a stable interlocking joint line across the foot. In a strict sense there are only three tarsometatarsal joints, as determined by the number of separate synovial cavities. The medial cuneiform first metatarsal joint is one. The intermediate and lateral cuneiforms and the second and third metatarsals form another, which contributes to the great tarsal synovial cavity. The third joint includes the cuboid and the fourth and fifth metatarsals. Therefore, we may refer to medial, intermediate, and lateral tarsometatarsal joints. Each of these three joints has its own fibrous articular capsule, parts of which are ligamentous. The dorsal and plantar ligaments are clearly capsular although some of their fibers may cross from joint to joint; the plantar are more variable. There is considerable variability among the dorsal, plantar, and interosseous tarsometatarsal ligaments (Fig. 1.4). Some are quite constant, many have variable attachments, others may be missing, and additional bundles may appear. The dorsal tarsometatarsal ligaments are short flat ribbons. They usually include one ligament from each metatarsal to its corresponding tarsal bone, plus two additional ligaments, from the second metatarsal base to the medial and lateral cuneiforms. The five relatively constant plantar tarsometatarsal ligaments are generally weak. The three bands attaching to the medial cuneiform are almost always present, but the ligaments on the lateral side of the foot are sometimes missing. However, they are reinforced by processes from the long plantar ligament and tibialis posterior tendon. Of the three interosseous tarsometatarsal ligaments, the medial one, also known as Lisfranc’s ligament, is the strongest and most important. It connects the lateral surface of the medial cuneiform to the medial surface of the second metatarsal base. Each of the three tarsometatarsal joint spaces has its own synovial membrane, and the spaces do not communicate with each other. However, they may communicate with other, adjacent synovial spaces. The largest of these spaces is known as the great tarsal synovial cavity. Centered on the intermediate tarsometatarsal joint, it includes the cuneonavicular, intercuneiform, cuboideonavicular, cuneocuboid, intermediate tarsometatarsal, and proximal intermetatarsal joints between the second and third, and third and fourth metatarsals.

1.2.10 Proximal intermetatarsal joints The three proximal intermetatarsal joints are plane synovial joints. The four lateral metatarsals are very firmly held in place at their proximal ends, partly through the strength of the ligaments which connect them to each other and to the tarsal bones, and partly by their interlocking arrangement. These joints are between the bases of the second and third metatarsals, the third and fourth metatarsals, and the fourth and fifth metatarsals. There is no joint between the bases of the first and second metatarsals and no ligaments connect them. There are, however, some weak connective tissue bands between these bones. The fibrous joint capsules attach to the bones superior, anterior, and inferior to the facets. The posterior edge of each metatarsal articular facet, except the first, blends with one of the tarsometatarsal joint surfaces so no capsule is present in those areas. There are dorsal, plantar, and interosseous ligaments connecting these four bones, all appearing to be thickenings of the fibrous capsules (Fig. 1.4). The interosseous ligaments are very strong, the dorsal are the weakest. The dorsal intermetatarsal ligaments are thin flat bands running transversely, that between the fourth and the fifth metatarsals being the strongest.

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The synovial spaces between the second and third, and the third and fourth metatarsals communicate with the great tarsal synovial cavity. The space between the fourth and fifth metatarsals is in communication with the lateral tarsometatarsal joint.

1.2.11 Distal intermetatarsal joints Traditionally, no joints are described among the heads of the metatarsal bones. However, they are bound by a series of short ligaments collectively called the deep transverse metatarsal ligament, which holds them together and helps limit movement among them. It is usually described as one long ligament with four parts. In that way, they resemble a series of syndesmoses. Each ligament connects the heads of neighboring metatarsals indirectly by attaching to the sides of the plantar metatarsophalangeal ligaments (Fig. 1.7).

1.2.12 Lesser metatarsophalangeal joints The metatarsophalangeal joints are all ellipsoid joints between the rounded heads of the metatarsals and the slightly concave articular surfaces on the bases of the corresponding proximal phalanges. The joints of all the toes are quite similar, but the hallux has certain specializations which merit its separate discussion in the next section. The second metatarsophalangeal joint is the most distal. Depending on foot size, the first and third are about 6 mm proximal to the second, and the fourth is an additional 6 mm proximal. The fifth is about 9 mm proximal to the fourth. Each joint has its own separate fibrous capsule, attaching to the metatarsals at their surgical necks. Dorsally the attachment is near the edge of the articular surface, but plantarly it is several millimeters proximal. The attachment to the phalanges is near the articular facets. The fibrous capsule is loose dorsally and plantarly, to allow for a wide range of motion. The plantar metatarsophalangeal ligaments are fibrocartilaginous pads which thicken the inferior parts of the capsules. They run from the neck of each metatarsal to the base of the proximal phalanx at the inferior edge of the posterior articular surface. They are firmly attached to the proximal phalanges, where they are thick and strong; but they are thin, flexible, and lax where they attach to the metatarsals. Therefore, the plantar ligaments move with the phalanges, and can even lose contact with the metatarsal heads in extreme flexion. Both surfaces of the plantar ligaments are smooth: the superior surfaces articulate with the heads of the metatarsals, while the inferior surfaces are grooved to provide pathways for the long flexor tendons. The smoothness of this ligament, along with the curve of the metatarsal heads, facilitates sideways motion at these joints. Collateral ligaments run obliquely anteriorly and inferiorly from the tubercles on either side of each metatarsal neck to the tubercles on the bases of the proximal phalanges. They are narrow, relatively strong bands. Fan shaped suspensory ligaments run between the metatarsal tubercles and the edges of the plantar ligament, just posterior to the collateral ligaments. Because of the loose fit of the fibrous capsular membranes, to allow flexion and extension, the synovial membranes are also loose. They are attached to the inner surfaces of the fibrous capsules.

1.2.13 Hallucal metatarsophalangeal joint The joint between the first metatarsal bone and the proximal phalanx of the hallux is, in its basic form, very similar to the joints of the lesser toes. However, it has undergone some modifications to accommodate two sesamoid bones embedded in the tendon of flexor hallucis brevis and to bear weight in certain phases of gait. The base of the proximal phalanx and the head of the first metatarsal have a fairly large congruent area. The articular facet on the phalangeal base is wider than it is high, and although it is roughly ovoid its bottom edge is almost straight. Its articular surface is smaller than the metatarsal head but forms a relatively deep concavity. There are two grooves running in an anterior-posterior direction on the plantar side of the head, separated by a distinct crest which lies lateral to the bone’s midline. These grooves hold the sesamoid bones of the joint; the medial is wider and deeper because the medial sesamoid is almost always the larger. There is some variation in the shape of the sesamoids, but they are generally ovoid. Their dorsal surfaces, which articulate with the first metatarsal, are slightly convex from side to side, concave from anterior to posterior, and covered with articular cartilage. The plantar surfaces are strongly convex. The sesamoids are embedded in the flexor hallucis brevis tendon, where they also receive fibers from the insertions of adductor and abductor hallucis.

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A fibrous capsule completely surrounds the joint. It is attached close to the edges of the articular surfaces dorsally but several millimeters away plantarly. The dorsal part of the capsule is thin; the plantar part is much thickened by fibrocartilage to form the plantar metatarsophalangeal ligament which, like its counterparts in the lesser toes, is a thick plate. Its attachment to the phalanx is firm while that to the metatarsal is loose. Therefore, the ligament’s position relative to the phalanx is relatively fixed but it may change its relationship to the metatarsal as the joint moves. The structure of this ligament is complicated by the presence of the sesamoid bones, their connecting ligaments, and the tendons of adductor and abductor hallucis. These structures form the sesamoid apparatus. The sesamoid bones are firmly embedded in the plantar metatarsophalangeal ligament, so they maintain a constant relationship to the proximal phalanx. The medial and lateral metatarsosesamoid ligaments are thickenings of the sides of the fibrous capsule, running from the tubercles on the metatarsal head to the sides of the plantar ligament, where they attach to the sesamoids. An intersesamoid ligament runs between the two bones. Additionally, two phalangeosesamoid ligaments connect the sesamoids to the proximal phalanx. These ligaments and the tendons attached to the sesamoid bones are all part of or blend with the plantar metatarsophalangeal ligament. The fibrous capsule is lined with a synovial membrane dorsally and at the sides. There are also pockets of synovial tissue on the plantar ligament, anterior to the sesamoids and near the metatarsal attachment.

1.2.14 Interphalangeal joints There are nine possible interphalangeal joints in each foot, one in the hallux and two in each of the lesser toes. However, fusion of the phalanges within a toe can occur, especially in the fifth, and in some populations the majority of people show such congenital symphalangy. Fusion of the middle and distal phalanges of the second, third, and fourth toes has also been reported but such cases are relatively rare. All the interphalangeal joints are ginglymus joints sharing a similar structure, although there are some small differences between the proximal and distal joints. The hallucal joint is considerably larger than any of the others. The heads of the proximal phalanges have definite sagittal grooves. The articular areas extend onto the plantar aspect for some distance but are scarcely visible on the dorsum. Small tubercles are located on the sides, dorsally. The heads of the middle phalanges are similar, except that the central grooves are much shallower, and there is only a slight extension of the articular facets plantarly. The bases of the middle and distal phalanges are approximately oval, with the long axes from running from side to side. All the bases are smaller than their mating phalangeal heads, only slightly concave, and divided in two by very low sagittal crests. The crests are lower on the distal phalanges. Each joint has its own capsule, thin dorsally where it is reinforced by the extensor tendons, and thicker on the sides and plantarly where it forms ligaments. The fibrous capsules attach to the bones a little distance from the articular facets, especially plantarly. Collateral ligaments are located on both sides of each joint. They extend anteroinferiorly from the tubercles on the head of the more posterior phalanx to those on the base of the more anterior. The sides of the capsules posterior to the collateral ligaments are strong, and help maintain the plantar ligaments. The plantar ligaments, like those of the metatarsophalangeal joints, are fibrocartilaginous. The superior surface of each ligament extends the socket on the phalangeal base, and the plantar surface provides a smooth grooved area against which the flexor tendon glides. The ligament is firmly connected to the phalangeal base, and so moves with it, but the connection to the head is loose and flexible. The plantar interphalangeal and metatarsophalangeal ligaments are converted into a continuous groove by thin layers of cartilage on the plantar sides of the bones. The synovial membrane lines each fibrous capsule. It forms a small pouch under the head of the more posterior phalanx, where the plantar ligament is loose.

1.3

Muscles and fascial specializations

A large number of muscles are involved with foot motion, some acting on the foot as a whole, others acting only within the foot, and still others involved both ways. Those which act on the foot as a whole—controlling motion relative to the leg—arise within the leg or distal part of the femur and travel into the foot where they insert. Some of them, such as digital flexors and extensors, also affect particular foot joints. The contractile portions of all these extrinsic muscles are located in the leg. Intrinsic muscles both arise and insert within the foot and therefore only affect motion at joints within the foot.

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1.3.1 Fascial specializations Fascial layers surround the leg and foot, separating the muscles from the skin. The superficial fascia is a loose connective tissue with varying amounts of adipose tissue within it. It is particularly thick in the sole. The deep fascia is a continuous sheet of membranous material surrounding the superficial faces of the muscles; it too is thickest in the sole, where it makes up the plantar aponeurosis. Additional fascial sheets extend inwards from this layer to attach directly or indirectly to bones. The spaces thus created compartmentalize the muscles into groups which tend to share similar functions and have similar origins, insertions, and neurovascular supplies. The anterior compartment of the leg contains muscles which are primarily extensors (dorsiflexors) of the ankle and extensors of the toes. The posterior compartment is subdivided into superficial and deep parts by the deep transverse intermuscular septum. The superficial part contains muscles which flex the knee and plantarflex the ankle. The deep part contains plantarflexors of the ankle, some of which also flex toes and invert or evert the foot. The lateral compartment muscles are plantarflexors and evertors. There are various ways of describing the compartmentalization of the sole, but they include at least three spaces: a medial compartment contains two intrinsic muscles acting on the hallux, a lateral compartment contains two muscles acting on the fifth toe, and an intermediate compartment containing a number of muscles acting mainly on the intermediate toes. Tendons of the extrinsic muscles also run through the sole. The layer of deep fascia is thicker in some regions than in others. Around the ankle some of these collagen fibers are oriented into retinacula which cover and form channels for the crural tendons passing deep to or through them, to ensure that they remain properly oriented relative to the joints they cross. The retinacula which guide the anterior crural tendons are located proximal and distal to the ankle joint. The superior extensor retinaculum runs between the anterior borders of the tibia and fibula, just proximal to the malleoli (Fig. 1.9). From medial to lateral, the muscles passing deep to it are tibialis anterior, extensor hallucis longus, extensor digitorum longus, and peroneus tertius. They are separated by two connective tissue septa; the last two muscles are in the same channel. The inferior extensor retinaculum is located on the dorsum of the foot. It is shaped like a sideways Y, with its stem on the lateral side. The stem splits into two parts, the superior attaching to the medial malleolus and the distal passing around the medial side of the foot to blend with the deep fascia of the sole. Extensor digitorum longus and peroneus tertius are held in place by the stem. Extensor hallucis longus and tibialis anterior are restrained by the medial bands. The flexor retinaculum covers the passageway connecting the deep part of the posterior compartment of the leg and the sole of the foot. It is a broad, roughly triangular thickening of the deep fascia, running between the medial edge of the medial malleolar sulcus and the medial surface of the calcaneus, guiding tibialis anterior, flexor digitorum longus, the posterior tibial vessels and tibial nerve, and flexor hallucis longus as they bend around the ankle and enter the sole. Strong septa create four channels, for each tendon and the neurovascular bundle. The tendons of peroneus longus and brevis travel deep to the superior and inferior fibular retinacula. The first of these runs posteroinferiorly from the lateral malleolus to the posterior end of the calcaneus. The second is continuous with the stem of the inferior extensor retinaculum, attaching inferiorly to the lateral surface of the calcaneus. They are roughly parallel. The plantar aponeurosis is a thickening of the deep fascia of the sole (Fig. 1.10). Two intermuscular septa separate the intrinsic plantar muscles into three compartments, medial, intermediate, and lateral. Their tension on the aponeurosis creates grooves in it, but subcutaneous fat fills the grooves so they are not visible on the skin.

1.3.2 Extrinsic dorsal muscles 1.3.2.1 Tibialis anterior Tibialis anterior arises from the lateral condyle and proximal half or more of the lateral surface of the tibial shaft, adjacent parts of the crural fascia and interosseous membrane, and fascia separating it from extensor digitorum longus. Its tendon of insertion begins about midleg, covering the deeper muscle fibers, then veers medially and passes through the medial compartment of the superior extensor retinaculum (Fig. 1.9). It then burrows through the superior arm of the inferior extensor retinaculum, where it is covered by only a few of the retinacular fibers, and deep to the inferior arm. As it runs along the medial side of the medial cuneiform it splits into two bundles, one inserting into the medial and plantar surfaces of that bone and the other, often larger, into the first metatarsal base. It acts primarily as a dorsiflexor of the foot and, because it inserts distal and inferior to the transverse tarsal joint, is also capable of inversion.

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FIGURE 1.9 Ankle retinacula and superficial structures of the foot, dorsal view.

1.3.2.2 Extensor digitorum longus The extensor digitorum longus belly lies medial to tibialis anterior. It has an extensive origin from the lateral tibial condyle, fibular head, two-thirds or more of the medial fibular surface, the proximal part of the interosseous membrane, and surrounding parts of the crural fascia, anterior intermuscular septum, and adjacent fascia. The tendon appears on the medial side of the belly about midleg, then descends almost vertically toward the foot. After passing deep to the superior extensor retinaculum it splits into four divisions, which cross deep to the stem of the inferior retinaculum and travel to the four lesser toes (Fig. 1.11).

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FIGURE 1.10 Plantar aponeurosis.

As they cross the proximal phalanges, the tendons of extensor digitorum brevis to the second through fourth toes are joined on their lateral sides by tendons of extensor digitorum brevis to form one tendon per toe. Each of these combined tendons, and the tendon to the fifth toe, divides into three parts near the proximal phalangeal heads. In each toe, the central part of the tendon inserts into the base of the middle phalanx dorsally. The medial and lateral slips reunite and insert into the base of the distal phalanx dorsally. The extensor digitorum longus tendons are major contributors to a membranous sheet which covers each of the four lateral toes from the metatarsophalangeal joint to beyond the distal interphalangeal joint. This sheet, the extensor

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FIGURE 1.11 Extensor tendons of the foot.

expansion, also receives contributions from other muscles and the deep fascia. It helps retain the extensor digitorum tendons in place and provides a means for them to extend the proximal phalanges. Circumferential fibers of the sheet at the level of the metatarsophalangeal joint are known as the extensor sling. Other fibers run from the anterior edge of the sling dorsally and distally, forming a small triangular sheet on each side of the central tendon and attaching to it, as far as the distal phalanx. These sheets are called the extensor wing or hood. Extensor digitorum longus dorsiflexes the ankle and extends the lesser toes as well as their interphalangeal joints.

1.3.2.3 Extensor hallucis longus Extensor hallucis longus arises from the middle of the medial surface of the fibula, next to extensor digitorum longus, and from the adjacent part of the crural interosseous membrane. The proximal part of the muscle is deep to tibialis anterior and extensor digitorum longus, but it becomes visible in the distal part of the leg shortly before it becomes

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tendinous. The tendon passes through the middle compartment of the superior extensor retinaculum, the superior limb of the inferior extensor retinaculum, and deep to the inferior limb (Fig. 1.11). Extensor hallucis longus inserts primarily into the base of the distal phalanx of the great toe, dorsally. A small accessory slip often branches from the medial side of the tendon proximal to the metatarsophalangeal joint and inserts into the joint capsule. An extensor expansion covers the dorsum of the great toe, similar to that seen in the lesser toes. The hallux and its interphalangeal joint are extended by this muscle.

1.3.2.4 Peroneus tertius Peroneus tertius is an extremely variable muscle, more or less fused with extensor digitorum longus. It arises from the inferior third or more of the medial fibular surface, just distal to the origin of extensor digitorum longus. It also receives fibers from the crural interosseous membrane and the anterior crural intermuscular septum. The belly is usually fused with extensor digitorum longus, even as far distally as the proximal part of its tendon. It lies along the lateral side of the long extensor, passing with it deep to the retinacula. The insertion is quite variable. It usually inserts into the dorsum of the base or the proximal part of the shaft of the fifth metatarsal, but it may also attach to the proximal end of the fourth metatarsal shaft or even send slips all the way to the fifth metatarsophalangeal joint. This weak muscle works with the other dorsiflexors at the ankle. It also likely helps the lateral leg muscles evert the foot.

1.3.3 Extrinsic plantar muscles The extrinsic plantar muscles arise proximal to the knee posteriorly and in the posterior compartment of the leg, where they are separated into two groups by the crural interosseous membrane. Gastrocnemius, soleus, and plantaris occupy the superficial part of the posterior compartment. They insert on the calcaneus and act on the foot as a whole. The remaining three posterior leg muscles arise within the deep part of the posterior crural compartment. They cross the ankle and various foot joints, thereby having more complicated sets of actions. The two digital flexors insert in the toes, but tibialis posterior has an extensive insertion onto all the tarsal and two or three metatarsal bones.

1.3.3.1 Triceps surae The component muscles of the triceps surae, soleus and the two-headed gastrocnemius, are almost entirely separate in their fleshy parts. However, they come together to form the strong calcaneal (Achilles) tendon, by which they act on the ankle joint (Fig. 1.12).

FIGURE 1.12 Muscles and tendons on the medial side of the foot.

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The medial head of gastrocnemius arises from the medial femoral condyle, superior to the articular area. The lateral head arises primarily from the lateral femoral epicondyle and from the posterior part of the lateral condyle. The heads remain separate until their insertion into an aponeurotic tendon near the middle of the leg. Soleus is wider and more massive than gastrocnemius. Its V-shaped origin lies along two bony lines: medially along the soleal line and middle third of the tibia’s medial border, and laterally from the head of the fibula and the posterolateral part of the posterior fibular surface. Additional fibers arise from a fibrous arch between the upper parts of the two bones. Connective tissue fibers appear on the muscle’s posterior surface, forming a tendon which becomes narrower and thicker until, about three fourths of the way down the leg, it joins with the gastrocnemius tendon. The calcaneal tendon is the thickest and strongest tendon of the body. The fibers spiral rope-like as they descend, those originating from gastrocnemius ending up on the lateral side and those from soleus medially. At the same time, the tendon becomes thicker and narrower until, just above the heel, it inserts into the posterior surface of the calcaneus. Both gastrocnemius and soleus are plantarflexors, the projection of the calcaneus posterior to the ankle giving them leverage. The oblique axis of the ankle joint also causes these muscles to act as invertors during plantarflexion. Both muscles contract more strongly when high heels are worn, apparently because of the instability resulting from a forward shift of the center of gravity. Gastrocnemius arises from the femur so it also flexes the knee, and is therefore less efficient as a plantarflexor unless the knee is extended.

1.3.3.2 Plantaris Plantaris is a muscle with a very small belly and a very long, very thin tendon. It arises from the femur near the lateral head of gastrocnemius, runs inferomedially between the two muscles of triceps surae, and descends to insert on or near the medial side of the calcaneal tendon. The length of the fleshy part ranges from about 4 11 cm. Compared to triceps surae, which does pretty much the same things it does, plantaris is a very weak muscle. It is a flexor of the knee and a plantarflexor of the ankle.

1.3.3.3 Flexor digitorum longus Flexor digitorum longus is the most medial of the muscles in the deep posterior compartment of the leg (Fig. 1.13). Its tendon continues into the sole, where it travels among the second layer of intrinsic muscles (Fig. 1.14). It gives origin or insertion to each of the intrinsic muscles in that layer. Its origin is from the posterior surface of the tibia, the septum which separates it from tibialis posterior, and from the deep transverse intermuscular septum. It runs nearly vertically down the leg, becoming tendinous and veering slightly to reach the medial side of the ankle. The tendon travels through the second compartment of the flexor retinaculum and lies in relation to the edge of the sustentaculum tali. In the sole of the foot it turns laterally and crosses plantar to the tendon of flexor hallucis longus, running between the first and third layers of intrinsic muscles. There, near the navicular, a bundle of fibers branches off the flexor hallucis longus tendon and joins the flexor digitorum longus tendon. The two long flexors are held against the navicular by a band of connective tissue, the “master knot of Henry.” The tendon of flexor digitorum longus then divides into four slips for the lesser toes. Just before this division it serves as the insertion for quadratus plantae, and as it divides its slips give origin to the lumbricals. Each slip lies immediately dorsal to a flexor digitorum brevis tendon, and from the level of the metatarsal necks the two flexor tendons to each toe travel within a common synovial sheath. When they reach the proximal phalanx each long flexor tendon passes through a slit in the brevis tendon, then inserts into the distal phalangeal base. Flexor digitorum longus is a plantarflexor and supinator of the foot, and a flexor of the metatarsophalangeal and interphalangeal joints of the four lateral toes.

1.3.3.4 Flexor hallucis longus Flexor hallucis longus arises from the distal two-thirds of the posterolateral part of the posterior surface of the fibula. It also takes origin from the deep transverse intermuscular septum, the posterior crural intermuscular septum, the distal part of the crural interosseous membrane, and the fascia which separates it from tibialis posterior. Its tendon of insertion begins within the muscle, just above its middle, and fleshy fibers continue to join the tendon as far as the ankle joint. It passes through the fourth compartment of the flexor retinaculum into the sole. There it runs directly anteriorly, inferior to the lateral head of flexor hallucis brevis, and when it approaches the metatarsophalangeal joint it shifts to pass between the hallucal sesamoids (Fig. 1.14).

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FIGURE 1.13 Extrinsic plantar muscle tendons in the foot.

Just after it enters the sole it is crossed on its inferior side by the tendon of flexor digitorum longus, and at that point it sends a variable fiber bundle to the digital tendon. In that way flexor hallucis longus can contribute to the function of the second through the fifth toes. Flexor hallucis longus inserts onto a rough V-shaped area on the plantar surface of the distal hallucal phalanx. The muscle can help plantarflex the ankle, and flexes the metatarsophalangeal and interphalangeal joints of the big toe.

1.3.3.5 Tibialis posterior Tibialis posterior lies deep in the posterior crural compartment, originating from almost the whole surface of the crural interosseous membrane. It also arises from the posterior surfaces of the tibia and the fibula, as well as the deep transverse intermuscular septum. Its contractile fibers converge upon a tendon which crosses deep to flexor digitorum longus and passes through the first compartment of the flexor retinaculum (Fig. 1.14). It inserts onto all of the tarsal bones except the talus, plus the bases of two or three metatarsals. As it enters the sole the tendon gives off a short strong band which attaches to the distal end of the sustentaculum tali, and the rest of the

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FIGURE 1.14 Extrinsic plantar tendons entering the sole.

tendon then divides into two unequal bundles. The smaller superficial part becomes flatter and wider, attaching to the tuberosity of the navicular and the inferior surface of the medial cuneiform. The deep part also flattens out, contributes to the origin of flexor hallucis brevis, then crosses the intermediate and lateral cuneiforms and the bases of the second and fourth metatarsals, and sometimes the base of the third metatarsal bone, to all of which it is attached. Tibialis posterior is an invertor and plantarflexor of the foot; it is generally described as the most powerful invertor.

1.3.4 Extrinsic lateral muscles The two muscles of the lateral compartment arise, as their names indicate, from the fibula. They are slender muscles which run almost straight down the leg, bend around the lateral malleolus, and insert in the foot.

1.3.4.1 Peroneus longus Peroneus longus arises from the head of the fibula, the superior half or two-thirds of the lateral surface of the fibula, the superior third of each crural intermuscular septum, and the proximal part of the crural fascia (Fig. 1.15). Its extensive origin usually also includes part of the lateral tibial condyle. The tendon of insertion begins high within the muscle. At

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FIGURE 1.15 Muscles and tendons on the lateral side of the ankle and foot.

the ankle it passes posterior to the lateral malleolus, but not directly in the lateral malleolar sulcus; peroneus brevis intervenes between it and the bone. It then changes direction to travel anteroinferiorly along the lateral surface of the calcaneus, inferior to the fibular trochlea, and as far as the fibular notch on the lateral surface of the cuboid. There it again changes direction and continues anteromedially across the sole of the foot. The main insertion of the peroneus longus tendon is the tuberosity on the lateral side of the first metatarsal base, plantarly. Smaller groups of fibers insert into the lateral side of the medial cuneiform and near the first metatarsal neck. This latter slip gives origin to much of the medial head of the first dorsal interosseous muscle. Peroneus longus is a plantarflexor, and because the last part of its tendon crosses the sole of the foot from lateral to medial, it is an evertor. It also helps maintain both the longitudinal and transverse plantar arches through the orientation of its tendon.

1.3.4.2 Peroneus brevis Peroneus brevis arises from the distal two-thirds of the lateral surface of the fibula, anterior to peroneus longus where their origins overlap, and the adjacent parts of the anterior and posterior crural intermuscular septa. The muscle fibers attach bipennately on a tendon which appears on its lateral side in the distal part of the leg, above the lateral malleolar sulcus. After passing around the ankle it runs anteroinferiorly across the calcaneus (Fig. 1.15). Peroneus brevis inserts onto the lateral part of the tuberosity of the fifth metatarsal, dorsally. It often sends a slip which inserts on one of the bones of the fifth toe or joins the long extensor tendon to the fifth toe. It is a plantarflexor and evertor of the foot.

1.3.5 Intrinsic dorsal foot muscles 1.3.5.1 Extensor hallucis brevis and extensor digitorum brevis Only one muscle mass is located on the dorsum of the foot (Fig. 1.16). Arising from the calcaneus it divides into four parts which travel to the hallux and to the second, third, and fourth toes. The part going to the hallux, which is often rather separate from the rest, is considered a muscle by itself named extensor hallucis brevis. The other three parts together make up extensor digitorum brevis. These short extensor muscles arise from the lateral part of the sulcus calcanei and from part of the lateral surface of the calcaneus. The muscle bellies lie deep to the tendons of peroneus tertius and extensor digitorum longus.

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FIGURE 1.16 Muscles, vessels, and tendons on the dorsum of the foot.

The long and short extensor tendons fuse in the other toes but they remain separate in the hallux. That of extensor hallucis brevis inserts into the base of the proximal phalanx of the great toe, dorsally. In the second, third, and fourth toes each combined long and short extensor tendon trifurcates. The central part reaches the base of the middle phalanx dorsally, while the lateral (mostly derived from extensor digitorum brevis) and medial slips recombine to insert on the base of the distal phalanx, dorsally. Extensor hallucis brevis extends the great toe. The other parts of the muscle extend the toes on the metatarsals and at the interphalangeal joints.

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1.3.6 Intrinsic plantar muscles There are eighteen intrinsic plantar muscles, muscles which both arise and insert within the sole. Starting from the plantar aponeurosis, the deep fascia of the sole, and moving inwards, they conveniently fall into four layers and will be described in this sequence: G

The most superficial (inferior) layer contains three muscles: abductor hallucis, flexor digitorum brevis, and abductor digiti minimi (Fig. 1.17).

FIGURE 1.17 Superficial plantar muscles.

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G

G G

PART | 1 Introduction

The second layer contains five distinct muscles, quadratus plantae and four lumbricals (Fig. 1.18). The tendons of flexors hallucis and digitorum longus travel along the foot in this layer. The third layer contains flexor hallucis brevis, flexor digiti minimi brevis, and the two headed adductor hallucis. The fourth layer, the deepest, contains the four dorsal and three plantar interosseous muscles. The tendons of peroneus longus and tibialis posterior also travel in this layer.

FIGURE 1.18 Superficial plantar muscles, with flexor digitorum brevis reflected.

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1.3.6.1 Abductor hallucis The soft muscle mass on the medial edge of the foot, bulging onto the sole, is formed by abductor hallucis (Fig. 1.17). It arises from the medial process of the calcaneal tuberosity and adjacent fasciae: the medial part of the plantar aponeurosis, the medial intermuscular septum, and the flexor retinaculum. At that level the retinaculum is split into deep and superficial layers, forming a sort of pocket in which abductor hallucis fits; the superficial layer of the pocket is continuous with the plantar aponeurosis. Abductor hallucis runs almost directly anteriorly along the medial edge of the foot. It covers the three tendons and the neurovascular bundle which enter the foot on the medial side, as they travel through the space known as the tarsal tunnel. More anteriorly it, or its long tendon of insertion, lies against the inferomedial aspect of the navicular, medial cuneiform, first metatarsal, and their joints. It partially covers flexor hallucis brevis, with whose tendon it blends. The insertion is largely into the metatarsophalangeal joint capsule, along with the medial tendon of flexor hallucis brevis. Part of the insertion may be on the medial side of the capsule, but it is mainly into the plantar ligament and medial sesamoid. The position of the tendon at the metatarsophalangeal joint makes it an abductor.

1.3.6.2 Flexor digitorum brevis Flexor digitorum brevis is a relatively large muscle, located just deep to the central part of the plantar aponeurosis (Fig. 1.17). It arises from the medial process and notch of the calcaneal tuberosity, the proximal third of the intermediate part of the plantar aponeurosis, and the two intermuscular septa of the sole. Its fibers run anteriorly, the belly becoming slightly wider as it approaches the metatarsal bases. There it divides into four bundles, each of which leads to a tendon for one of the four lateral toes. The two medial bundles are the largest and the one to the fifth toe is the smallest. The tendons continue anteriorly, lying plantar to the tendons of flexor digitorum longus, and enter the fibrous flexor sheaths of the toes. Just after entering the fibrous flexor sheath, near the metatarsophalangeal joint, each tendon splits longitudinally. The halves then come together again, leaving a sort of slit through which a flexor digitorum longus tendon passes to become more superficial. Each brevis tendon then divides again, with some of its fibers decussating, near the head of the proximal phalanx. The halves insert into the shaft of the middle phalanx on its sides, plantarly. This muscle plantarflexes the metatarsophalangeal and proximal interphalangeal joints.

1.3.6.3 Abductor digiti minimi Abductor digiti minimi is a long slender muscle, contributing to the soft tissue mass on the lateral edge of the foot (Fig. 1.17). The origin of abductor digiti minimi on the plantar surface of the calcaneus extends from the lateral process of the tuberosity to the medial process, where it is anterior to the origin of flexor digitorum brevis. The muscle also takes origin from the lateral intermuscular septum and the lateral part of the plantar aponeurosis. A variable amount of tissue joins its deep surface from the shaft of the fifth metatarsal. The muscle travels anteriorly along the lateral border of the foot, near the sole. The tendon of insertion forms along its superficial side near the calcaneocuboid joint, receiving fleshy fibers as far distally as the metatarsophalangeal joint, becoming thick and strong. The tendon inserts into the lateral side of the fifth metatarsal head, the joint capsule, and the lateral side of the proximal phalangeal base, plantarly. Although capable of abducting the little toe, its plantar location probably renders it more effective as a flexor.

1.3.6.4 Quadratus plantae Quadratus plantae, also known as flexor accessorius, has two heads of origin which come together as they approach their insertion (Fig. 1.19). The medial head arises from the medial surface of the calcaneus, posteriorly. The lateral head, smaller than the medial, arises from the plantar surface of the calcaneus anterior to the lateral process, and may also arise from the lateral side of the long plantar ligament. The two heads come together to form a sheet-like, V-shaped muscle. The muscle fibers run almost directly anteriorly until they meet the flexor digitorum longus tendon, where they insert. The muscle fibers of the two heads have somewhat different insertions. The medial head joins the deep side of the long flexor tendon proximal to its division, and contributes to the tendons of the second, third, and usually fourth toes. The lateral head usually penetrates the long flexor tendon proximal to its division and joins it from within. By applying traction to the flexor digitorum longus tendon, quadratus plantae flexes the lesser toes as does the long flexor, acting on the metatarsophalangeal and interphalangeal joints.

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FIGURE 1.19 Structures deep to the flexor digitorum brevis.

1.3.6.5 Lumbricals The four lumbricals are small “worm-shaped” muscles, arising from soft tissue and inserting almost entirely into soft tissue. The first lumbrical is the most medial. It acts on the second toe, not the first, and so on. For about half the length of each division of the flexor digitorum longus tendon, lumbrical fibers arise from it (Fig. 1.20). The first lumbrical arises from the medial side of the most medial division of the flexor digitorum longus

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FIGURE 1.20 Relations of the long flexor tendons.

tendon. The rest arise between the divisions of the tendon just as they separate: the second arises from the first and second divisions, the third from the second and third divisions, and the fourth from the third and fourth divisions. All except the first lumbrical are bipennate. The first lumbrical is the largest, with the others about equal in size. The lumbricals insert into the medial sides of the extensor expansions on the second, third, fourth, and fifth toes. These muscles cross plantar to the metatarsophalangeal joints and, through their attachment to the extensor expansion, dorsal to the interphalangeal joints. Thus they flex the metatarsophalangeal and extend the interphalangeal joints. Since they travel on the medial sides of the toes, they move their toes medially at the metatarsophalangeal joints. The horizontal motions at those joints are referred to the resting axis of the second toe, so therefore the first lumbrical is said medially abduct the second toe and the others adduct their toes.

1.3.6.6 Flexor hallucis brevis Flexor hallucis brevis lies plantar and parallel to the first metatarsal (Fig. 1.21). It originates mainly by a flat Y-shaped tendon, whose medial arm is actually a slip from the tibialis posterior tendon of insertion. The lateral arm of the tendon runs roughly transversely, and is anchored mainly by separate slips to the plantar aspects of the cuboid and lateral cuneiform.

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FIGURE 1.21 Muscular relationships inferior to the metatarsophalangeal joints.

The muscle belly begins near the medial intercuneiform joint and runs distally along the lateral border and inferior surface of the first metatarsal, where it develops a complex insertion. Near the head of that bone the belly splits into two parts which go toward the sides of the first metatarsophalangeal joint, become tendinous, and insert onto the proximal phalanx. The lateral part is joined by the tendon of adductor hallucis and the medial part by the tendon of abductor hallucis. Fibers of the plantar metatarsophalangeal ligament merge with the conjoined tendons. The combined structure, formed by contributions from three muscles and the plantar ligament of the metatarsophalangeal joint, is sometimes referred to as the plantar plate. This complex contains two sesamoid bones, one in each side, which articulate with the metatarsal head. This muscle is a flexor of the metatarsophalangeal joint of the great toe.

1.3.6.7 Flexor digiti minimi brevis Flexor digiti minimi brevis is a small muscle which lies inferior to the fifth metatarsal (Fig. 1.22). It has a fleshy origin from the base of that metatarsal (sometimes extending onto the shaft), the fibular sheath, and the promontory of the cuboid. The belly runs parallel to the fifth metatarsal toward the metatarsophalangeal joint. It joins abductor digiti minimi, which lies plantar to it, to insert together into the head of the fifth metatarsal, the metatarsophalangeal joint capsule, and the base of the proximal phalanx laterally and plantarly. The muscle is a flexor of the fifth metatarsophalangeal joint.

1.3.6.8 Adductor hallucis Adductor hallucis is made up of two distinct bellies which come together near their insertion (Fig. 1.22). The oblique head arises from the fibular sheath in the region of the tarsometatarsal joints, and from the bases of the second, third, and fourth metatarsals. It runs anteromedially toward its insertion, lying just lateral to flexor hallucis brevis.

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FIGURE 1.22 Plantar interossei and related structures.

The smaller transverse head arises from the fibrous capsules and plantar metatarsophalangeal ligaments of the third, fourth, and fifth metatarsophalangeal joints and from the nearby parts of the deep transverse metatarsal ligament, not directly from bone. It crosses the foot on the plantar side of the deep transverse metatarsal ligament, and bends dorsally between the first two metatarsals to join the lateral edge of the oblique head, close to its insertion. The combined muscle blends with the medial head of flexor hallucis brevis, and inserts onto the lateral sesamoid and the lateral side of the base of the proximal phalanx. Both parts of adductor hallucis are flexors of the hallux, although the angle of the transverse head relative to the joint limits its effectiveness. Similarly, both heads are positioned to adduct the proximal phalanx, but in this case the transverse head is more advantageously placed.

1.3.6.9 Dorsal and plantar interossei The interossei are a group of muscles lying among the metatarsals (Fig. 1.22). They fill the intermetatarsal spaces, bulging slightly onto the dorsum of the foot and overlapping the metatarsal shafts plantarly.

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They can be divided into two groups with slightly different functions and relations. The four dorsal interossei are the larger, filling most of each intermetatarsal space and bulging above and below the bones. The three plantar interossei are smaller, occupying the inferior parts of their intermetatarsal spaces. Altogether, they form an almost continuous sheet of muscle on the plantar side of the metatarsals. Because of the way the interossei lie in the intermetatarsal spaces and the nature of their action it is necessary to redefine the terms adduction and abduction for them. The plantar interossei are all placed so as to bring their toes toward the resting axis of the second toe while the dorsal interossei move their toes away from that axis. The resting axis of the second ray, not the midline of the body, is thus the reference line for describing the action of these muscles. The Plantar muscles ADduct their toes—bring them toward the reference line—and the Dorsal ABduct them—bend them away from that axis—in a transverse plane. The dorsal interossei are numbered one through four from medial to lateral. They are bipennate, arising from the two bones bounding each intermetatarsal space, with their fibers converging upon a common tendon. Additionally, they take origin from the ligaments of the tarsometatarsal joints, the fibular sheath, and the fascia of the adjacent muscles. Usually the first also has a significant origin from a slip of the peroneus longus tendon. Their tendons of insertion pass dorsal to the deep transverse metatarsal ligament to attach to both sides of the proximal phalanx of the second metatarsophalangeal joint, and the lateral sides of the third and fourth joint capsules. They also join the extensor expansions of the toes. They are plantarflexors and abductors of the second (medially and laterally), third, and fourth toes. The plantar interossei are unipennate, arising from the medial surfaces of the third, fourth, and fifth metatarsals, and thus lie in the three lateral intermetatarsal spaces. Their soft tissue attachments are similar to those of the dorsal muscles, but only in involving one metatarsal. They cross dorsal to the deep transverse metatarsal ligaments and insert on the medial sides of the proximal phalanges, joint capsules, and extensor expansions. They are plantarflexors and adductors of the third, fourth, and fifth toes.

1.4

Nerves

With the exception of the saphenous nerve, a branch of the femoral nerve, all nerves reaching the foot are derived from the sacral plexus; but not all nerves formed in that plexus supply the foot. In fact, the nerves which are sensory to the foot and motor to the muscles which act upon and in it are all traceable to only two derivatives of the plexus, the tibial (L4,5, S1,2,3) and common fibular (L4,5, S1,2) nerves. These two nerves travel next to each other in a common fibrous sheath, where they form what is called the sciatic nerve, from the sacral plexus until they reach about two-thirds of the way down the posterior of the thigh. There they separate. The sciatic looks like one nerve but its two components are separated by fascia the whole length of the time they travel together, and there is no interchange of axons between them.

1.4.1 The fibular nerves in the foot The common fibular nerve continues inferiorly, winds around the neck of the fibula, descends in the anterior compartment of the leg, and soon after splits into the deep and superficial fibular nerves (Fig. 1.23). Before it divides, however, it gives off the sural communicating branch, which will contribute to the formation of the sural nerve. The deep fibular nerve continues to travel inferiorly within the anterior compartment of the leg. It passes deep to the superior extensor retinaculum, then crosses the ankle to enter the foot. At that point it is accompanied by the dorsal artery of the foot. It soon divides into lateral and medial terminal branches. The lateral terminal branch of the deep fibular nerve runs distally and laterally across the dorsal surface of the foot, sending branches to extensors digitorum and hallucis brevis. It has a number of small branches which innervate the ankle and intertarsal joints, and the tarsometatarsal, metatarsophalangeal, and interphalangeal joints of the second, third and fourth toes, and the skin of those regions. The medial terminal branch accompanies the dorsal artery of the foot, giving off articular branches to the more medial intertarsal joints, a branch to the first metatarsophalangeal joint, and sometimes a muscular twig to the first dorsal interosseous muscle. When it reaches the first intermetatarsal space the medial terminal branch divides into two dorsal digital nerves, which supply the adjacent sides of the hallux and second toe. The superficial fibular nerve usually travels in the lateral compartment of the leg, but it may pierce the anterior crural intermuscular septum to descend within the anterior compartment. In the distal third of the leg it becomes superficial, enters the foot, and provides most of the cutaneous innervation of the dorsum. In the most common pattern it soon divides into its terminal branches, the medial and intermediate dorsal cutaneous nerves of the foot. The medial dorsal cutaneous nerve passes superficial to the superior extensor retinaculum to supply the skin of the foot, and at the ankle

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FIGURE 1.23 Superficial structures on the dorsum of the foot.

joint it divides into two dorsal digital nerves. The medial or proper dorsal digital nerve to the hallux supplies the medial side of the great toe. The lateral branch innervates the adjacent sides of the second and third toes. The intermediate dorsal cutaneous nerve runs across the foot toward the fourth toe, where it divides into two common dorsal digital nerves supplying the adjacent sides of the third and fourth, and the fourth and fifth toes. It also supplies skin on the lateral side of the dorsum of the foot and communicates with the lateral dorsal cutaneous nerve.

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1.4.2 The tibial nerves in the foot The tibial nerve is the larger component of the sciatic nerve continues inferiorly down the thigh, through the popliteal fossa and the deep posterior compartment of the leg, accompanied by the posterior tibial vessels. At a variable level in the upper part of the leg it gives off the medial sural cutaneous nerve, which becomes superficial and joins the lateral sural communicating branch of the common fibular nerve to form the sural nerve. This nerve reaches the foot, where it sends branches to the ankle and subtalar joints and to the skin around the lateral malleolus. Its name then changes to the lateral dorsal cutaneous nerve, whose branches supply the lateral side of the foot and the fifth toe. The trunk of the tibial nerve passes through the third compartment of the flexor retinaculum. While in that space in that space it gives off articular branches to the ankle and cutaneous branches to the posteromedial part of the sole, then splits into the medial and lateral plantar nerves (Fig. 1.24). The medial plantar is the larger of these two nerves. After leaving the space deep to the retinaculum, it runs between abductor hallucis and flexor digitorum brevis, supplying them and the first lumbrical. Articular branches supply nearby joints of the tarsus and metatarsus, especially the talonavicular, cuneonavicular, and intercuneiform joints. The remainder of the nerve then terminates by dividing into common plantar digital nerves, which reach the skin of the three medial interspaces. These nerves serve not only the plantar skin, but also nearby metatarsophalangeal and interphalangeal joints and the skin around and deep to the distal half or more of the medial three and a half nails. The lateral plantar nerve travels laterally, passes through the medial intermuscular septum, and continues between flexor digitorum brevis and quadratus plantae (Fig. 1.25). Articular twigs supply the calcaneocuboid joint and other branches supply the anterolateral part of the foot. The trunk of the nerve supplies quadratus plantae and abductor digiti

FIGURE 1.24 Neurovascular structures passing through the tarsal tunnel.

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FIGURE 1.25 Neurovascular structures in the plantar midfoot.

minimi. Then it divides into superficial and deep branches. The superficial branch innervates the fourth interspace and the lateral side of the foot, plantarly. There are articular twigs to the regional joints and muscular ones to flexor digiti minimi brevis and the third plantar and fourth dorsal interossei. The deep branch innervates both heads of adductor hallucis, all the interossei except those of the fourth intermetatarsal space, and the lateral three lumbricals. It sends small branches to the nearby intertarsal and tarsometatarsal joints. It has no cutaneous components.

1.5

Blood supply

1.5.1 Arteries The main supply of blood to the foot is through three large vessels: the anterior tibial, posterior tibial, and fibular arteries. The first two are the terminal branches of the popliteal artery high in the posterior of the leg, and the fibular artery is a major branch of the posterior tibial. The anterior tibial artery quickly passes through the crural interosseous membrane into the anterior compartment of the leg, where it descends in an almost straight line toward the ankle joint and from then on is known as the dorsal artery of the foot (arteria dorsalis pedis). The posterior tibial artery veers medially as it travels inferiorly, enters the third compartment under the flexor retinaculum and, under cover of abductor hallucis, bifurcates into the medial and lateral plantar arteries. The fibular artery arises high in the back of the leg, veers laterally, and parallels the fibula until it crosses the ankle joint and breaks up into several

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posterior lateral malleolar arteries. Additionally, medial and lateral sural branches off the popliteal artery nourish the superior parts of gastrocnemius and soleus. The arterial branches in and around the foot and their anastomoses are numerous and quite variable; only some of the most important are mentioned here. This helps maintain the circulation as pressures from standing and moving constrict individual vessels. The dorsal artery of the foot continues across the talus, navicular, intermediate cuneiform, and base of the second metatarsal to the first intermetatarsal space, where it divides in two. Its branches along the way include the lateral tarsal artery and two medial tarsal arteries in the area of the navicular, and the arcuate artery at the base of the second metatarsal. The arcuate artery runs laterally across the metatarsal bases, where it gives off the second, third, and fourth dorsal metatarsal branches and then communicates with the lateral tarsal artery. The dorsal metatarsal arteries end near the bases of the toes by each splitting into dorsal digital arteries. The terminal branches of the dorsal artery of the foot are the deep plantar branch, which dives inferiorly between the first and second metatarsals to anastomose with arteries in the sole, and the first dorsal metatarsal artery, which nourishes the first intermetatarsal space and the hallux. The posterior tibial artery divides, deep to the flexor retinaculum, into the medial and lateral plantar arteries. Just before this it gives off two or three medial calcaneal branches which nourish the calcaneal tendon and the inferomedial side of the heel. The medial plantar artery travels within the medial compartment of the foot supplying skin, abductor hallucis, flexor digitorum brevis, and the first dorsal interosseous muscles. It also helps supply the medial side of the great toe, and superficial branches join the medial three plantar metatarsal arteries. The larger lateral plantar artery runs anterolaterally toward the fifth metatarsal base, where it bends, becomes the plantar arch, and runs medially back across the metatarsal bases. It ends by anastomosing with the deep plantar branch. Along the way, it gives off four plantar metatarsal arteries, which anastomose with their dorsal equivalents, and each of them ends by splitting into two plantar digital arteries. Within each toe, the plantar digital arteries form a tuft of blood vessels which underlie and extend up the sides of the distal phalanx.

1.5.2 Veins There are two sets of veins in the lower limb, superficial and deep, named according to their relationship to the deep fascia. The superficial veins form a rete in the superficial fascia, communicate with the deep vessels, and usually do not accompany named arteries. Deep veins accompany the arteries, usually in pairs called venae comitantes (singular, vena comitans). Communicating or perforating veins tie the two parts of the system together. In the foot valves, which usually direct blood from superficial to deep vessels, are frequently missing.

1.5.2.1 Superficial veins The dorsal digital veins on adjacent sides of neighboring toes come together to form dorsal metatarsal veins. These drain into to the dorsal venous arch, near the metatarsal bases. At its lateral end the arch leads into the lateral marginal vein. The medial dorsal digital vein of the hallux joins the medial end of the arch to form the medial marginal vein. At the ankle these veins become the small and great saphenous veins, respectively. These veins all receive contributions from the dorsal venous network, which covers the dorsum of the foot. Note that there is only one set of dorsal veins in the toes, and that they drain mainly into the superficial system even though they accompany arteries. On the plantar aspect the digital veins lead into the deep venous system. However, there is a plantar venous network, which communicates both with the deep veins and the marginal veins of the dorsum.

1.5.2.2 Deep veins Plantar digital veins unite to form plantar metatarsal veins, after sending branches to the dorsal surface. Here, as dorsally, the metatarsal veins drain into a venous arch, but the plantar venous arch is deep and accompanies the plantar arterial arch. The ends of this arch are the medial and lateral plantar veins. From this point onwards their paths are similar to those of the arteries in reverse, receiving tributaries comparable to arterial branches. The deep veins on the dorsum carry relatively little blood, but they follow the arteries.

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1.5.3 Lymphatics The superficial lymph vessels fall into two sets. The medial lymphatics, draining the dorsum and medial side of the foot, approach the great saphenous vein and, in a general way, follow it. The lateral lymphatics, which drain the lateral side of the foot, mostly lead into lateral vessels. Deeper lymph vessels follow the paths of the main arteries proximally. There are no lymph nodes in the foot. Arterial and Neural Supply to Foot Muscles Muscle Arteries Nerves Extrinsic Dorsal Muscles Tibialis Anterior

Anterior Recurrent Tibial

Deep Fibular; L4,5

Genicular Anastomosis

Deep Fibular; L5, S1

Anterior Tibial

Deep Fibular; L5, S1

Anterior Tibial

Deep Fibular; L5, S1

Gastrocnemius

Medial and Lateral Sural

Tibial; L5, S1

Soleus

Medial and Lateral Sural

Tibial; L5, S1

Plantaris

Sural

Tibial S1,2

Flexor Digitorum Longus

Posterior Tibial

Tibial; L5, S1,2

Flexor Hallucis Longus

Posterior Tibial

Tibial; L5, S1,2

Posterior Tibial

Tibial; L4,5

Fibular

Superficial Fibular; L5, S1

Fibular

Superficial Fibular; L5, S1

Dorsal Artery of the Foot

Deep Fibular; L5, S1

Dorsal Artery of the Foot

Deep Fibular; L5, S1

Abductor Hallucis

Medial Plantar

Medial Plantar; S1

Flexor Digitorum Brevis

Medial Plantar

Medial Plantar; S1,2

Genicular Anastomosis Anterior Tibial Extensor Digitorum Longus Anterior Tibial Fibular Extensor Hallucis Longus Fibular Peroneus Tertius Extrinsic Plantar Muscles

Fibular Posterior Tibial

Fibular Tibialis Posterior Fibular Extrinsic Lateral Muscles Peroneus Longus Anterior Tibial Peroneus Brevis Anterior Tibial Intrinsic Dorsal Muscles Extensor Hallucis Brevis Lateral Tarsal Extensor Digitorum Brevis Lateral Tarsal Intrinsic Plantar Muscles

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Abductor Digiti Minimi

Lateral Plantar

Lateral Plantar; S2,3

Quadratus Plantae

Lateral Plantar

Lateral Plantar; S2,3

Lumbrical 1

Lateral Plantar

Medial Plantar; S2,3

Lumbricals 2, 3, 4

Lateral Plantar

Lateral Plantar; S2,3

Flexor Hallucis Brevis

Lateral Plantar

Medial Plantar; S1,2

Flexor Digiti Minimi Brevis

Lateral Plantar

Lateral Plantar; S2,3

Adductor Hallucis

Lateral Plantar

Lateral Plantar; S2,3

Dorsal Interosseous 1

Medial Plantar

Lateral Plantar; S2,3

Dorsal Interosseous 2

Lateral Plantar

Lateral Plantar; S2,3

Lateral Plantar

Lateral Plantar; S2,3

Lateral Plantar

Lateral Plantar; S2,3

Lateral Plantar

Lateral Plantar; S2,3

Second Dorsal Metatarsal Dorsal Interosseous 3 Third Dorsal Metatarsal Dorsal Interosseous 4 Fourth Dorsal Metatarsal Plantar Interossei 1, 2, 3

Further reading Dubois JF, Levame JH. Anatomie descriptive du pied humain. Paris: Librairie Maloine; 1966. Kelikian AS. Sarrafian’s anatomy of the foot and ankle. 3rd ed. Philadelphia: Wolters KluwertLippincott Williams & Wilkins; 2011. Standring, S. Gray’s anatomy, 4th ed. Churchill Livingstone Elsevier, 2008. Paulsen F., Waschke J. Sobotta atlas of human anatomy. In: General anatomy and musculoskeletal system 1.

Chapter 2

Basic Biomechanics Joseph M. Iaquinto1,2,3 1

RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 2Department of

Mechanical Engineering, University of Washington, Seattle, WA, United States, 3Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Abstract An understanding of tissue, joint, and limb function requires a basic understanding of biomechanics. Biomechanics is the study of movement in living organisms, and its features include the interaction of forces, moments, mechanical properties and more, on the motion or equilibrium of body segments. Considering the lower extremity, the movement of the foot through a volume of space can be achieved by the coordinated activation of muscles in the thigh, leg, and foot—and also from contributions of the entire body moving through space. Gait is a clear example of a human body advancing forward in space with rhythmic harmony between bones, muscles, tendons, cartilage, and ligaments. This chapter briefly reviews some core mathematical and conceptual concepts associated with biomechanics and guides the reader toward rich sources of deeper knowledge on these topics. The intent of this chapter is to provide additional context for the reader regarding the entire content of this book.

2.1

Introduction

The concepts of biomechanics are typically taught from a mechanical engineering perspective to make them more relatable and to simplify them before layering on the additional complexity that comes with biological systems. Depending on the background of the reader, phrases like “the knee or elbow joint is a hinge” or “muscles pull on bones” or “cartilage cushions your joints” may be familiar, and while they provide basic context to help their audience relate to simpler mechanical systems they may interact with every day (a door, a seatbelt, a pillow), they only scratch the surface on what is truly occurring in these tissues and tissue systems (the elbow is actually a collection of joints, none of which act along either a single or simple axis like a hinge; muscles do contract, but can impart stabilizing, dislocating and rotating actions on the bones they anchor in; cartilage cushioning is an understatement to the complex and multi-phasic behavior this tissue utilizes to prevent wear while being dynamically burdened with forces well over bodyweight, thousands of times a day). Slightly more accurate and more complex concepts may quickly earn their keep in providing the reader with a better perspective on the motions and forces that they personally experience every day. This chapter covers the very basic concepts in biomechanics, and attempts to bring additional context of the following chapters.

2.2

Terminology

A brief formalization of terms and their definitions can serve as a reference point. 1. A force is a phenomenon that can act on a body to change its state of motion. Sir Issac Newton’s second law of motion describes the relationship between force (F), mass (m) and acceleration (a) [Eq. (2.1)] [1]. F 5 ma

(2.1)

2. Related to force is stress, which is the action of a force with respect to a given area. For example, an object experiencing a stress of 1 Newton per meter squared (1 N/m2) is experiencing 1 Newton of force, for every square meter of area. 3. If that force is acting perpendicular to the surface of the object, it is noted as normal stress. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00031-7 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 2.1 Talus, lateral view. Moment of rotation (green arrow) is generated about the center of mass of the bone due to the application of a force (blue arrow) on the talar head, applied perpendicular and at a distance (blue dotted line) from the talar center of mass.

4. If that force is acting parallel to the surface of the object, it is shear stress. 5. A moment is force that is acting at a distance to a point on an object. In biomechanics, moments tend to act on the state of rotation of an object by imparting a torque (Fig. 2.1). 6. The change in shape of an object (typically due to the application of force) is deformation. 7. Strain is an expression of the deformation of an object from its original geometry. 8. Similar to force, a strain acting perpendicular to a cross section of material is noted as normal strain. 9. A strain acting parallel to a cross section of material is noted as shear strain. 10. Equilibrium is a state whereby the sum of the forces acting on an object is zero, typically this is expressed for both forces and moments [Eqs. (2.2), (2.3)]. X X

Fðx; y; zÞ 5 0

(2.2)

Mðx; y; zÞ 5 0

(2.3)

There are several terms relevant to the field of biomechanics, some of which are included in the topics in the following sections. For a deeper reading on these terms, the author recommends several entry level and more complex texts on the topic [25].

2.3

Statics

Statics is the study of external forces acting on a system, which is in a state of equilibrium. Consider the situations where a human body is in a state of equilibrium, such as quietly standing, or poised on a step from one level to the next, but not moving, or balanced briefly on the tips of the toes. In these cases, the object (the body, or the lower extremity) is not moving as the external effects of gravity, and the internal generation and application of force by the muscles is exactly balanced to achieve that equilibrium. In calculations for statics, in addition to the state of equilibrium, there must also be an assumption that the bodies which are at rest are considered rigid (nondeformable). This is an ideal time to introduce a very useful concept in the field of biomechanics (which is also useful generally in the field of mechanics!)—the free-body diagram. Free-body diagrams are a mechanism that can be used to show objects of interest, and the magnitudes and directions of forces acting on those objects. In biomechanical analysis, the pathway to generating a free-body diagram to accurately represent an anatomical analysis can be challenging. Doing so requires an understanding and appreciation of the anatomical area of interest, including the shape of bones, the insertion sites for tendons or muscle origins, the location and shape of bony articulations, and any additional tissue (such as ligaments) that may act on and constrain the system of interest. A further requirement is that the anatomy is well enough understood to make justifiable assumptions or simplification of the anatomy to both sufficiently answer the question of interest, and to allow the problem to be sufficiently constrained as to be solvable. Example 2.1: A shopper is reaching for an overhead container, the shelf that the container is resting on is just out of reach, so the shopper raises up on their toes to gain enough extra height. Wary of upsetting the contents of the shelf, the shopper pauses on their toes, as they use their arms and hands to secure the container. Our interest in this scenario is to analyze the forces acting on joints or locations within their foot while they were briefly in that state of equilibrium (Fig. 2.22.4).

FIGURE 2.2 Anatomical illustration of the foot during rise onto toes. Achilles tendon (blue) and gastrocnemius (red). While only containing a fraction of the true anatomical detail present in the foot, this diagram illustrates how immediately complex this anatomy is, thus providing evidence for the utility of a simplification to a free-body diagram for high level consideration of this mechanical environment.

FIGURE 2.3 Simplification of the anatomy into an initial free body diagram. Hind and midfoot bones have been fused into a single unit with endpoints at the Achilles tendon insertion, the tibiotalar articulation, and the first metatarsophalangeal joint (1st MPTJ). Arrows representing the lines of action of the Achilles tendon (red, Fach), tibial joint reaction (green, Ftt), and ground reaction at the 1st MTPJ (purple, Fgrd).

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FIGURE 2.4 More complete free-body diagram, showing decomposition of the achilles, tibiotalar, and ground reaction loads into their respective x and y components, as well as segment length information. theta, (Θ) denoting foot segment angle to ground, and distances from the ground to the tibiotalar joint (dtt) and Achilles tendon (dach). A diagram such as this readily lends itself to a calculation of forces and moments (in this case both are constrained to zero as this is a static example). The objective of this process is to identify assumptions (segment rigidity within the foot, and only considering one muscular element) to allow for simplification of modeling the load state (between the ground, ankle joint, and Achilles tendon).

2.4

Dynamics

Dynamics is the study of external forces acting on an object, which results in a change of motion. Consider the propulsion of a person up a stair step, or the cessation of motion from stepping off a curb, or the rhythm inherent in gait, creating numerous differing instances of dynamic motion throughout, which is referred to as the gait cycle. In these instances, the activation of muscle forces and the poise of the body segments are acting to generate motion or to cease motion—the goal is a change from one dynamic state to another, or into or out of a state of equilibrium. As with statics, dynamic calculations make the assumption that the bodies are rigid, but unlike statics, the net forces acting on a system are not zero (not in equilibrium). Dynamics can be further parsed into two subfields. Kinematics is the study of motion, independent of the forces that cause it. The variables of interest in this field tend to be time, position, velocity, acceleration, and geometry, but not forces, stresses, deformation, or strains. Various motion capturing systems (discussed elsewhere in this text) are designed to accurately and precisely capture these motions using various visible light, magnetic or X-ray sensor technologies. Kinetics is the study of motion and the forces and moments which cause it. The variables of interest here expand to include forces and moments. Incorporating force sensing elements into a measurement system that already has motion capture capability (such as a gait laboratory with force plates) enables the measurement of enough information to perform kinetic analysis of human motion. On the topic of human motion, kinesiology is a related term here, which is the study specifically of human movement and motion.

2.5

Strength of materials and deformation

The examples provided previously all contain the assumption of rigid body motion, and this is of course never the case in human biomechanics—but this assumption is both practical and useful for many types of human biomechanic analyses. For many other types of analysis, consideration must be made for the deformation of objects which are experiencing forces. As examples, consider that most human tissues, especially structural tissues, respond (strengthen, maintain, or weaken) based on the mechanical stimulation those tissues are under. This translates to stresses and strains felt by the cells in those tissues, and the responses such perturbations trigger within the cells and in the extracellular matrix. Bone and muscle for example, will strengthen over time to accommodate high loads, and will resorb or atrophy over

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FIGURE 2.5 Idealized representation of a biological tissue during a load to failure test. Stress versus strain plot showing tissue behavior during loading. Linear region (where elastic modulus calculation is derived) is shown as a blue bar. Yield point is indicated by the blue dot. Ultimate stress is indicated by the green dot. Yield is indicated by the red dot.

time in the absence of mechanical load. Adipose tissue, in addition to shielding neurovascular channels within the body, also plays a structurally cushioning role in the foot, and particularly in the plantar fat pad of the heel. Some additional core concepts relating to strength of materials and deformables are defined in the following: 1. Elastic modulus—is the ratio of stress to strain in a materials region of elastic response, elastic modulus is a material property 2. Stiffness—is the resistance to deformation of an object, stiffness is a structural property 3. Failure—is a complex term with numerous criteria in engineering, but for the purposes of biomechanics, consider it catastrophic breakdown of a material (a bone breaking or a tendon tearing) 4. Yield/ultimate/failure load and strain—these are specific points along a material load to failure curve that represent the end of the elastic response (yield), the end maximum load or strain the material experienced (ultimate), and the point of catastrophic breakdown (failure) (Fig. 2.5). 5. Poisson’s ratio—the ratio of transverse strain in a material, to the axial strain in a material (consider a tendon becoming thinner under axial load). The concepts of deformation and strain become the key in this field of study. Like other concepts, we can start with very clean and simplified examples, and work toward the more complex reality. There are two major types of deformation commonly discussed in biomechanics: elastic deformation and plastic deformation. Elastic deformation occurs when a material (such as a tissue) is perturbed by load or strain, causing the material to deform from its original shape. When that perturbation is removed, the material will return to its original shape. Plastic deformation is seen when after the perturbation is removed, the material does not return to its original shape and is permanently changed by the perturbation. Various materials experience drastically different amounts of elastic and plastic deformation. Bone has very little elastic or plastic deformation and will rapidly progress to failure/fracture as it deforms. Skin on the other hand has a comparably massive range of elastic deformation before injury. Another type of deformation is fluid flow. The human body and most body tissues have a very high fluid content, and the solid and fluid portions of human tissues can interact in fascinating ways to give rise to very unique mechanical behavior. This brings us to the last major concept of this chapter.

2.6

Viscoelasticity

There are substantial differences between engineered materials and biological tissues. Engineered materials tend to be homogenous, experience very small deformations (for typical structural engineered materials), and exhibit a significant linear range of behavior. Biological tissues do not typically align with any of those behaviors. Biological tissues are

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FIGURE 2.6 Spring and dashpot model elements. Spring (left) icon and equation, force in the spring is a function of the spring constant (k) and the _ strain (x) in the spring. Dashpot (right) icon and equation, force in the dashpot is a function of the coefficient of viscosity (µ) and the rate of strain (x) applied.

FIGURE 2.7 Maxwell, Kevin-Voight, and Standard Linear Solid Models for viscoelasticity. Arrangement of the model elements dictates the overall behavior of the model, selection, and matching of spring constants and viscosity constants modulates the behavior further. These models have both served directly in the literature and as starting points for more complex model formulation.

complex blends of liquid and solid elements, acting to allow for a huge range of human motion and substantial deformations. The fluid content of biological tissues gives rise to another feature of this tissues, inherent viscosity, or resistance of that fluid to shear motion. Viscoelasticity is a combined consideration of the solid and fluid properties of biological tissues. This tends to provide for tissues which are both strain and stain rate dependent, which deform under the application of load and experience recovery, but not necessarily instantaneous recovery. As with most complex behaviors, mathematical models serve to unify how we can compare and think about viscoelastic materials. Three basic models, the Kelvin-Voigt, Maxwell, and Standard Linear Solid models, comprise of combinations of two key elements, a spring (representing the solid component of the tissue) and a dashpot (representing the fluid component of a tissue) (Fig. 2.6). The behaviors of these three models are described (Fig. 2.7). These models and other models for viscoelasticity have been applied to many tissues of the body including ligaments [6], organ tissues [7], bone [8], cartilage [9], and skin [10]—to name a few.

2.7

Summary

As noted, the intent of this chapter is not to provide a comprehensive background on biomechanics, but rather to call out some key concepts and bring them to the readers’ attention. Concepts of loads versus deformation, the common viscoelastic properties and behavior of tissues, and the ability to decompose a complex anatomical environment into a more relatable interaction of joint reaction forces, muscle forces and external loads, will hopefully serve the reader in the rest of this text to understand the benefits of the technologies employed to evaluate foot biomechanics, the pathologies that plague this anatomy, and approaches to their repair and restoration. It is strongly recommended that the reader obtains and consults a dedicated biomechanics reference text if they pursue deeper knowledge of any of the topic areas presented in this text.

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References [1] [2] [3] [4] [5] [6] [7] [8] [9] [10]

Principia. | Library of Congress. ,https://www.loc.gov/item/2021667054/. London: Royal Society, 1687. [accessed 30.3.22]. Levangie PK, Norkin CC, Lewek MD. Joint structure & function: a comprehensive analysis. 6th edition, 2019, 535. Knudson D. Fundamentals of biomechanics. Fundam Biomech 2021;. ¨ zkaya N, Goldsheyder D, Nordin M, Leger D. Fundamentals of biomechanics: equilibrium, motion, and deformation, 4th ed. 2016 Jan 1. O p. 1454. Fung Y-C. Biomechanics. 1993 ,http://link.springer.com/10.1007/978-1-4757-2257-4. [accessed 30.3.22]. Pen˜a E, Pen˜a JA, Doblare´ M. On modelling nonlinear viscoelastic effects in ligaments. J Biomech 2008;41(12):265966. Wang X, Schoen JA, Rentschler ME. A quantitative comparison of soft tissue compressive viscoelastic model accuracy. J Mech Behav Biomed Mater 2013;20:12636. Johnson TPM, Socrate S, Boyce MC. A viscoelastic, viscoplastic model of cortical bone valid at low and high strain rates. Acta Biomater [Internet]. 2010 6(10):40734080. ,https://pubmed.ncbi.nlm.nih.gov/20417735/. [accessed 30.3.22]. Hislop BD, Heveran CM, June RK. Development and analytical validation of a finite element model of fluid transport through osteochondral tissue. J Biomech [Internet]. 2021 123. ,https://pubmed.ncbi.nlm.nih.gov/34048964/. [accessed 30.3.22]. Al-Benna S. Establishing tension-free direct wound closure using the viscoelastic properties of the skin. J Cutan Med Surg [Internet]. 2014 18(5):307315. ,https://journals.sagepub.com/doi/10.2310/7750.2013.13137?url_ver 5 Z39.88-2003&rfr_id 5 ori% 3Arid%3Acrossref.org& rfr_dat 5 cr_pub110pubmed. [accessed 30.3.22].

Chapter 3

Anatomical Nomenclature: Conundrums of Nonstandardized Foot and Ankle Terminology Thomas M. Greiner Department of Health Professions, University of Wisconsin—La Crosse, La Crosse, WI, United States

Abstract The terminology of the foot and ankle can be confusing. The fact that several different terms can be associated with each structure or concept complicates their understanding. Many practitioners may assume that terminological differences are associated with specific disciplines, such as clinician versus bioengineer, or basic versus applied researcher. However, that is not the case always. Terminology differences are more complicated. As such, opportunities for collaborations among laboratories, disciplines, languages, and cultures are inadvertently stifled. Different organizations have attempted to address this problem by developing recommended terminological standards. A few of these standards are regularly consulted, but none have been fully successful. This chapter outlines some of the more common areas of terminological confusion that may be found in discussions of foot and ankle biomechanics. Some confusions arise in the use of unfamiliar words, such as eponyms or different regional names. These words are presumed to provide concise and simple descriptions. But they can sometimes have the opposite effect. Other conceptual terms are common in their use, but complex in their meaning. Important concepts, such as the anatomical position or the anatomical planes, are not associated with simple and consistent meanings. The effect of differing concept meanings can cascade through the understanding of other terms that, on the surface, would have uncomplicated meanings and applications. This problem applies particularly to descriptions of movements, which can be misunderstood because they are related to the irregularly established concepts of the anatomical planes. Finally, there are terms of motion that can have identical definitions, but when their underlying mathematical assumptions differ, they can be associated with widely contrasting values. Some suggestions are offered to minimize miscommunication within scientific discourse. In the end, however, we must remain on guard for terminology problems. By understanding where, and how, terminological differences can intrude into presentations, we can try to mitigate their influence on the interpretation of results.

3.1

Introduction

You keep using that word. I do not think it means what you think it means. Inigo Montoya, The Princess Bride [Motion Picture]. United States: Act III Communications.

Terminology is one of the first lessons in professional training. We come away from this lesson believing that we are now equipped with a lexicon that will be generally understood by all with training in related disciplines without regard to language, culture, or academic interest. For the most part this impression is valid. There are a great many concepts that can be successfully communicated across many potential intellectual barriers by employing our professional terms. However, as our research emphasis becomes more specialized (such as the specifics of the foot and ankle instead of the generalized organism) our applications become more diverse (for example clinical, commercial, industrial, comparative, and/or evolutionary development). We may, therefore, find that our presumably universal terminology sometimes fails to transcend one or more intellectual barrier. Many attempts have been made to address this problem by creating standardized terminology lists. One of the first attempts at standardizing anatomy terminology came from the 1895 Kongress der Anatomischen Gesellschaft in Basel, Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00005-6 © 2023 Elsevier Inc. All rights reserved.

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Switzerland. The result of this congress became known as the Basle Nomina Anatomica. While this document set the pattern for future standardization efforts it did not gain worldwide acceptance. Subsequent attempts to update and adopt this anatomical standard came in 1933, 1936, 1955, 1966, 1977, 1980, and in its current form as Terminologica Anatomica [1] (TA) for human anatomy in 1997 and as Nomina Anatomica Veterinaria [2] (NAV) for nonhuman anatomy in 1999. While other terminology standardization documents associated with clinical and professional organizations also now exist, these two documents have the best current claims for a standardization of anatomical terminology. Although competing terminological standards are in general agreement, they are never identical. Some words are left off each list, and frequently lists favor different quasi-synonyms; words that mean basically the same thing, but convey a slightly different emphasis or are associated with very different secondary meanings. Although many word lists are constructed to be discipline centric, the crossdisciplinary nature of most scientific research can make these focused standards counter-productive to intellectual advancement. In addition, terminology standards are rarely associated with definitions. Therefore, the standard cannot insure that words are being used to convey the same meaning by all of its users. The more focused our intellectual applications, the more likely it becomes that we use terminologies that are not equally understood outside of our own laboratory or clinic. A natural expression of scientific hubris will, from time to time, cause us all to assume that we are the ones who are using the appropriate standardized terminology. When differences are found it must be the other researcher who has erred. Research [3 6] shows that this assumption is not always valid. Despite the well intentioned efforts of many professional organizations it is a sad fact that carefully constructed lists of standardized terminologies are rarely followed, enforced, or even consulted. Therefore, this chapter will not attempt to create a new standard of correct terminologies or overtly recommend one list over another. Instead, it will point out some of the more common sources of terminological confusion. In some cases existing standards will be offered as the best possible terminological choice. In other cases, no best terminological choice seems to exist. The aim, then, is to suggest greater care in our descriptions so that we do not rely upon terminologies that may not be understood as we intended them to be.

3.2

Anatomical descriptions

3.2.1 The persistence of eponyms Ideally anatomical terminology would be based on a unifying logic—a description of shape, location, or function. Were this logic in place an unfamiliar term would be decipherable and an unfamiliar structure or function would be nameable. Unfortunately, no such system exists. A single descriptive perspective cannot always be applied, especially in the face of long standing tradition. In some cases tradition will continue to supersede any attempt at applying a unifying logic. An example of this is the use of eponyms. An eponym is a term that is based on the name of its discoverer or for a person who has otherwise become associated with the structure or concept. When something is truly novel, naming it after someone is akin to including a reference to the original description. Still, the general consensus within terminology standards is to discourage the use of eponyms because they lack descriptive value, are frequently culturally and linguistically specific, are often associated with multiple terms, and sometimes honor persons whom current mores would prefer not to recognize [1,2,7 9]. Even so, there is a strong contingent that values the continued and expanded use of eponyms [8,10 12] because their use serves as a symbol of membership within the community. Certainly our colleagues in the physical sciences have no concerns about using eponyms such as Boyle’s or Ohm’s Laws. Where these disciplines intersect with our own, we are unlikely to abandon terms such as Euler and Cardan angles or Newtons of force. Similarly, a plethora of eponyms exist within diagnostic medicine, even when the terms are unintentionally misleading, such as Charcot-Marie-Tooth disease. While a sense of identity within a professional community certainly has value, when it goes too far it has the effect of exclusion. This diminishes opportunities for cross-disciplinary, and cross-cultural, research. Evidence from publication behaviors [12] suggests that it would be naive to believe that eponyms will ever disappear from biomedical discourse. Nonetheless, scientific discourse would probably benefit from the limited use of eponyms, to be replaced by linguistically and culturally neutral alternatives whenever possible.

3.2.2 Regional descriptions 3.2.2.1 The distal element of the lower limb Eponyms are not the only obscure terms that are used within professional groups. Many biomechanists and clinicians, for example, demonstrate a preference for the word “shank” to indicate the distal portion of the pelvic limb. This term

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is often preferred because “leg” can refer to the entire limb and not just to the distal segment. “Shank” has a long history in English, being derived from Old English and Middle English equivalents for “leg” [13]. The preference for “shank” seems to rely on its unfamiliarity in modern English. Even still, “shank” does not correspond with only one anatomical definition. In English “shank” can convey meanings that are completely unrelated to the foot and ankle, which adds to the difficulties of interpretation for non-English speaking researchers. The relevant anatomical connotations also include the entire pelvic limb, all or part of the pectoral limb, and can be restricted to mean the proximal segment of the pelvic limb (the thigh) [13]. An added complication is that “shank” also references a part of the shoe [13]. When researchers are interested in the effects of shoe design on limb mechanics, we might find ourselves discussing the influence of the shank upon the shank. Because of these many definitions, “shank” displays no superior communicative clarity and may actually be counterproductive. Inasmuch as they are the official references on anatomical nomenclature, TA and NAV provide the Latin word crus to describe the portion of the limb found between the knee and the ankle. Crus translates into English as “leg.” Continuing with their suggestions, the portion of the limb between the knee and the hip is femur, which translates as thigh. The entire limb is TA: membrum inferius or NAV: membrum pelvinum, which in English is the lower or pelvic member. It is important to note that when these Latin terms correspond with common English terms; “leg” can be restricted to mean knee-to-ankle even in common English [13]. If the recommended anatomical nomenclature were used exclusively, and consistently, there would be no need for an alternative term. Still, even if the ambiguous meaning of “leg” persists so that readers are unable to resolve it to mean either the entire limb or merely its distal segment, in nearly all discussions of the foot and ankle the distinction between those two meanings is irrelevant. Attempts to improve clarity by using an alternative term is unnecessary.

3.2.2.2 Forefoot and hindfoot Functional investigations frequently divide the foot into two or three units. The words “forefoot” and “midfoot” appear to be uncontroversial in these applications, although they do have problems. Naming the proximal part of the foot as either hindfoot or rearfoot is the focus of some controversy. As heated, and inconsequential, as the argument between those two choices can get, both sides seem to miss an important problem. The terms forefoot and hindfoot-rearfoot do not incorporate broader anatomical considerations. In comparative contexts, the word “forefoot” might be confused with the forward appendage (TA and NAV provide manus, which translates to hand) and the hindfoot-rearfoot term could mean the foot as a whole (officially pes, which translates as foot). In common usage the “forefoot,” or “forward foot,” is synonymous with hand, particularly when discussing nonprimates. Many clinicians might dismiss this particular conundrum as irrelevant, due to their preoccupation with the human animal. But, from a scientific perspective, there is an opportunity for greater clarity. Available terminology provides the terms proximal and distal to describe positions along the length of a limb. Greater clarity might be obtained by using the terms distalis pede (distal foot) and proximalis pede (proximal foot), instead of forefoot and hindfoot, at least once within the manuscript and/or as keywords.

3.2.3 Applying the anatomical position The anatomical position is drilled into us during our earliest training. We are comfortable with the idea that this position provides the standard reference against which all limb postures, planes of perspective, and movements are described. Without this common reference we would have no ability to accurately discuss and compare our biomechanical analyses or clinical techniques. It may, therefore, surprise many readers to learn that there is no standardized anatomical position that has obtained official sanction from the larger professional and scientific organizations. The term, and the posture it describes, has been defined by tradition and general purpose rather than by any consideration of how it is applied in scientific discourse. Many classic anatomy textbooks assume knowledge of the anatomical position, but never describe it (e.g., Gray’s Anatomy [14] and Cunningham’s [15]). Neither TA nor NAV include “anatomical position” as an official term. Most definitions of the human anatomical position (HAP) take the form: the subject stands erect and facing forward, upper limbs by the side with the palms facing forward. Great care is mentioned about the orientation of the hands, so that it creates a posture where the bones of the forearm are parallel. Such care may not seem necessary for the pelvic limb, for in human anatomy the two leg bones are always roughly parallel. Perhaps for this reason the position and orientation of the feet are often neglected. When foot position is mentioned, we are provided with statements such as “feet together,” [16] “heels together and the toes pointing somewhat outward” [17], “feet facing forward” [18], or “feet parallel” [19].

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Often a picture is supplied where the orientation of the feet do not match well with the written descriptions (see for example the 2005 edition of Gray’s Anatomy [18]), or there are several pictures that seem to contradict each other (see for example Moore, et al. 2014 [19]). These descriptions and illustrations can be difficult to interpret. We are left with varying foot orientations that shift the meaning of some basic anatomical directions as applied to the foot. The questions that arise (Fig. 3.1) are: is the foot positioned to be roughly parallel to the leg or does it lie in a roughly perpendicular orientation? And, are the feet aligned roughly parallel to each other or are they positioned so that the toes diverge? The differences in these foot orientations are essential when investigating some important methodological, theoretical, and evolutionary implications for understanding how the human foot and ankle functions. Within the Tetrapoda, the superclass of vertebrate animals with a homologous foot and ankle anatomy, there are three basic foot postures: unguligrade, digitigrade, and plantigrade (Fig. 3.2). Unguligrade animals, such as horses and cows, walk on their phalanges usually with the aid of a hoof. The foot bones of this type of animal are typically arranged in a column that can be seen as an extension of the leg and thigh. The ballerina standing en pointe demonstrates a human approximation of this position. In humans this posture can be seen as the leg parallel foot anatomical position (FAP). Digitigrade animals, such as birds, dogs, and cats, walk on their metatarsal heads and the heel process does not contact the ground. Humans practicing so called “toe walking” approach this foot posture. Plantigrade foot postures are most similar to the ancestral foot orientation (Fig. 3.3). This posture is characterized by metatarsals that lie roughly parallel to the ground. The plantigrade posture characterizes animals such as amphibians, reptiles, monotremes, some rodents, most primates, and closely resembles the foot and limb posture of the tetrapod embryo. A fully plantigrade posture, as practiced by humans and other apes, is characterized by a calcaneal tuberosity (heel) that also contacts the ground. The foot position of the tetrapod embryo differs slightly in that the limbs have not undergone developmental rotation. Because of this, the toes are oriented laterally, with the first digit (hallux and pollex) on the anterior-cranial side. This first digit position is often described as “pre-axial” while the fifth digit is “post-axial.” In these descriptions “axial” references the developmental axis of the foot. By considering the three basic foot postures (Fig. 3.4), the HAP that requires a roughly leg parallel foot orientation makes sense as a basis for comparative analyses with digitigrade and unguligrade forms. This is especially true since most plantigrade animals can approximate this leg parallel position, while adopting the plantigrade foot posture would be difficult for the other forms. Added to this complexity is that humans are the only animals that habitually stand with knee and hip joints at near full extension. A flexed hip and knee posture characterizes all other plantigrade animals, whether those animals are standing on all fours or are temporarily standing bipedally. The flexed hip and knee posture means that the standing ankle in these animals is held in dorsiflexion. Only humans habitually stand with the foot perpendicular to the leg. So, the “neutrally oriented,” perpendicular, foot is uniquely human. Analyzing subjects from this referent posture only makes sense in the realm of human clinical applications. The large accumulation of clinical data that makes use of this HAP is difficult to reconcile with the needs of comparative or developmental questions.

3.2.4 Application specific human anatomical positions The definition of the HAP is vague concerning the proper orientation of the foot. Sometimes this lack of specificity can provide some advantages. The anatomical position with a loosely defined, or completely undefined, foot orientation is what would most likely be used in biomechanical studies where actions of the foot and ankle are being collected, compared, and analyzed along with actions of the knee, hip and perhaps other joints. Because a specific orientation of the foot is not part of the posture’s definition, researchers can evaluate the effects of variables like femoral torsion, tibial torsion, hip rotation, knee rotation, and other factors that may produce in-toeing or out-toeing in the overall gait patterns. It would be difficult to compare these types of data when the postural definition specifies the orientation of the foot. Imagine how an identical foot position could be described as displaying in-toed orientation when the referent position is defined with “toes pointing somewhat outward” [17] or as showing an out-toed orientation when the referent position is defined with “toes directed anteriorly” [19]. Although the HAP is the fall back referent for all clinical and biomechanical studies, it is actually not often applied in studies that focus on the foot and ankle. More focused investigations make use of a Leg Anatomical Position (LAP) that is based on a consideration of landmarks located solely within the leg and foot. Frequently this is constructed as a local reference frame. The referent posture for the foot and ankle recommended by the International Society of Biomechanics (ISB) [20], which makes use of the medial and lateral malleoli and the medial and lateral tibial condyles, serves as the most common example of an LAP. Within this position the orientation of the foot can be evaluated because it is not part of the definition—note that evaluation of tibial torsion is specifically addressed in the ISB standard. Still, foot orientation based on femoral torsion, or hip or knee rotation cannot be addressed. In this system,

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FIGURE 3.1 The anatomical directions as conventionally applied to the body. When applied to the foot, the meaning of anatomical directions will change based on the assumed orientation of the foot in anatomical position. The top two images compare directions when the toes are oriented to point forward (top left) or point outward (top right). Some of the foot specific meanings of these two orientations are illustrated with the middle two images, where the vectors of foot medial lateral can change their relationships to the vectors of body medial lateral. The bottom two images compare directions along the sagittal plane when the foot is in a leg parallel position (bottom left) or a leg perpendicular position (bottom right). All of these perspectives can be found in association with the definition of the anatomical position. The leg perpendicular posture seems most intuitive for students of human anatomy and biomechanics, and it is the standardized posture employed within this book. However, this foot posture is not common in nature, and is even impossible to achieve among many mammals. The anatomical position is described with a different suite of foot orientations for many members of the chordate phylum, and for humans during embryological development.

anterior posterior is defined as the cross product of the medial lateral and inferior superior vectors. This direction is frequently called the “floating” axis because its direction relative to the foot can change from one subject to the next. Whether or not they apply the ISB standard, a great many biomechanical assessments make use of a LAP in that they

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FIGURE 3.2 The three basic foot postures within the tetrapoda (plantigrade, digitigrade, and unguligrade) and their relationships to standardized anatomical planes. In the plantigrade posture as practiced by most humans (top image) the coronal and transverse planes of the leg correspond with those found in the torso. However, these planes can be 90 degrees off when applied to the foot. In the typical digitigrade posture the coronal and transverse planes of the leg and foot are roughly perpendicular to those of the torso. In addition, the leg and foot planes are slightly different from each other. In the unguligrade posture the leg and foot coronal and transverse planes are again roughly perpendicular to their torso orientations, although the leg and foot are more nearly equal. The human foot to leg perpendicular posture, which is the standard orientation utilized within this book, is actually the more unusual posture among the tetrapoda. As such, that posture has limited value for comparative and evolutionary investigations of the foot.

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FIGURE 3.3 The application of anatomical directions to basic body forms. The top image presents the primitive or ancestral tetrapodal condition, where the proximal limb segment sprawls laterally from the torso, the leg segment orients 90 degrees inferiorly, and the foot reorients another 90 degrees into a plantigrade posture. In this position, anterior posterior and cranial-caudal describe the same directions, while dorsal ventral and superior inferior are the same. This body form can be seen as a slight modification of the embryological condition (central image), the difference here being with the rostral-caudal (cranial-caudal) axis that bends with the body. Both of these body positions are contrasted with directions in the basic mammalian form (bottom image).

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FIGURE 3.4 Comparison of anatomical directions as applied to the foot and leg in three primate (human order) configurations. In all images, anterior posterior and superior inferior are shown as they would apply to the torso in anatomical position. The top images shows the limb and foot in the leg perpendicular position. Note that the orientations of dorsal ventral and proximal-distal are perpendicular to those described for the foot. The bottom image presents anatomical directions as applied in the leg parallel foot position. Here dorsal ventral is roughly parallel to anterior posterior and proximal-distal is roughly parallel to superior inferior. It is also noteworthy that the directions in the foot are close to parallel to those of the leg and body. The central image represents the plantigrade posture that is typical of most nonhuman primates. While this posture is not typically associated with any concept of the HAP, it is an important concept for primate comparisons and for human evolutionary studies.

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do not record, or report, anatomical orientations of the opposite limb or of the body as a whole. This is the beginning of a methodological blind spot that focuses too much on the foot and ankle and does not consider the interactions between the foot and the rest of the body. The researchers, and clinicians who conduct and apply this research, seem to assume that the LAP and the HAP are the same, or very similar. That assumption is not always justified. This methodological blind spot is taken further when researchers employ a FAP by limiting the orientation consideration solely to landmarks located within the foot. These studies are completely incapable of describing the orientation of the foot relative to the body, or of one foot relative to the other. They typically make up for this shortcoming by focusing on intrinsic foot structures. The foundational work by Inman [21,22] is an example of this type of work. Indeed, Inman’s work has set many of the clinical standards for the “normal” orientations and functions of the intrinsic foot joints, especially for the subtalar joint. Yet, because the referent orientation contains no information about the body as a whole there is no proper way to convert these clinical norms into values relative to a whole body anatomical position. Problems with the anatomical position concept are not special to the foot and ankle. For example, standardized positioning of the shoulder is notoriously difficult, and even the thorax and abdomen will change relative orientations due to the breathing cycle. In many applications, an anatomical position without critical appraisal is more than adequate. However, scientific discourse requires a more precise framework so that we can focus on differences of interpretation rather than differences of expression. Unfortunately, it would seem that some investigations of the foot and ankle may be reporting different findings primarily because of different assumptions about the anatomical position and its terminological application. The impact of the differing anatomical positions will become even more apparent as the nonstandard meanings of other terms are discussed.

3.2.5 Defining anatomical directions, planes, and axes 3.2.5.1 Posterior anterior versus ventral dorsal By strict definition, anterior means “in the front” or “in the direction of forward travel” [13,23]; posterior is the opposite. The word ventral is derived from the Latin word ventralis, and relates to the abdomen or belly. Ventral means “of the belly” or “in the direction of the belly.” Dorsal, also a Latin derivative, means “of the back” or “in the direction of the back.” In most applications of the HAP (Fig. 3.1) these terminological couplets are treated as synonyms and they are presented as such in many beginning anatomy textbooks (e.g., Drake, et al., 2012 [24] and Moore, et al., 2015 [25]). Treating these terms as synonyms is officially sanctioned within by TA [1], even though this official list is inconsistent in treating these terms as synonyms. These term couplets are not synonymous in most non-HAPs, and this is one stated reason why NAV was created as a separate document [2]. More importantly, in the ancestral/embryonic posture dorsal ventral pair with superior inferior and posterior anterior pair with caudal-cranial (Figs. 3.2 and 3.3). Conventional anatomical terminology as applied to the pelvic limb reflects the embryonic posture prior to developmental limb rotation. Pelvic limb surfaces and muscular compartments that are anatomically anterior are also identified as dorsal structures. If the human adopts a leg parallel foot position all dorsal surfaces of the foot and leg will be directed anteriorly. When the HAP is defined with the foot perpendicular to the leg, the now superior foot surface retains its dorsal designation; the sole of the foot is the ventral surface. This confusion of what is posterior anterior and ventral dorsal is not special to the foot and ankle; similar terminological redirections can be found in descriptions of the brain and the heart. While these terminological redirections do have a logical basis in their reference to the embryonic position, the choice of when and how to retain embryonic positions is not always clear. These redirections can be seen as examples of applied traditions that are not always easy to justify.

3.2.5.2 Anatomical planes Standard anatomical description includes three cardinal planes that by all definitions must be mutually orthogonal. By convention these planes are defined based on their application to the torso and that application is extended into the limbs. The median plane (or median sagittal plane) describes the plane that divides the torso into equal left and right parts. Because the limbs are off the center line, this median application never applies to studies of the foot and ankle. Instead, sagittal planes (also known as the paramedian or parasagittal planes) parallel the median plane. Depending upon the orientation of the foot in the HAP, a sagittal plane may divide the foot obliquely. In applications of the LAP and FAP the sagittal plane often starts as the plane that best approximates a medial lateral divide through the leg or foot, and all other sagittal planes lie parallel to this reference. The foot (and leg) median plane now references the midline of the foot. This midline may have a functional, rather than geometric, definition and therefore foot median need

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not create an equal medial lateral divide. Because of the separate referent bases, the leg and foot sagittal plane need not be parallel to that of the body. Nonetheless, construction of a sagittal plane, either overtly or assumed, serves as the referent. All other cardinal planes must be orthogonal to this plane and to each other. The other two cardinal planes are the coronal plane (also known as the frontal plane) and the transverse plane (also known as the horizontal, axial, or transaxial plane). These two planes will exchange how they divide the torso, and by extension the limb and foot, based on how the planes are defined, how these definitions are applied to the different anatomical positions, and how they are applied to the different segments of the leg and foot. The coronal plane can be defined in two ways, as separating (A) anterior from posterior, or as separating (B) ventral from dorsal. The transverse plane has three working definitions; it separates (1) superior from inferior, (2) cranial from caudal, or when applied to the limbs, it separates (3) proximal from distal (Fig. 3.5). In the HAP, and in the nonhuman bipedal position, for most of the leg the two definitions of the coronal plane accomplish the same thing even though the posterior anterior and dorsal ventral vectors are pointed in opposite directions. If continued from its torso definition, the coronal plane will pass obliquely through the limbs that stand in a flexed posture. The three definitions of the transverse plane are roughly similar with the exception that definitions 1 and 2 may section the limb somewhat obliquely in the nonhuman biped, while definition 3 would have the plane run perpendicular to the long axis of the limb. In the quadrupedal/ancestral/embryonic anatomical position the definitions of these two planes apply somewhat differently and these applications change based on a whole body versus limb restricted perspective. In the torso of a quadruped, the coronal plane that splits ventral from dorsal (definition A) would end up dividing proximal from distal in the limb (transverse plane definition 3). Similarly, the transverse plane that separates cranial from caudal in the quadrupedal position (definition 2) would split anterior from posterior (dorsal from ventral) in the limb (coronal plane definitions A and B). These applications can be reversed if alternative definitions are used. The quadrupedal coronal plane that follows definition A and splits anterior from posterior (a very unusual definition when applied to the quadrupedal torso) would extend into the limb and continue to separate anterior from posterior. The quadrupedal transverse plane definition 1 that divides superior from inferior (again a very unusual application to the quadrupedal torso) would extend into the limb and continue to divide superior from inferior (proximal from distal). A slightly different set of circumstances will obtain when these planes are applied to the foot. In the leg parallel HAP applications to the foot are only slightly oblique to their applications within the leg. The coronal plane continues to divide dorsal from ventral and anterior from posterior. In this posture, the transverse plane in the foot also roughly follows all three definitions; separating (1) superior from inferior, (2) cranial from caudal, and (3) proximal from distal. In the digitigrade, and unguligrade quadrupedal postures a similar continuity with the leg applies. However, these planes change dramatically when the foot is considered in a plantigrade (perpendicular) position. In the foot perpendicular HAP, the coronal plane that lies parallel to the plane as defined for the torso and leg divides the foot anterior posterior, following definition A, but does not follow definition B and instead divides proximal from distal (transverse plane definition 3). In this posture the transverse plane established for the leg and torso divides the foot superior from inferior (definition 1) but does not follow definitions 2 and 3, and instead divides ventral from dorsal (coronal plane definition B). The foot perpendicular bipedal posture transposes two of the plane definitions. In the quadrupedal plantigrade/ancestral/embryonic posture, the coronal and transverse planes as defined within the torso apply well to the foot according to coronal definition B and transverse definitions 2 and 3. Coronal definition A and transverse definition 1 may also apply, depending upon how the toes are positioned within this posture. There is no way to simultaneously apply all of the definitions for the coronal and transverse planes. Many clinicians and researchers who do not focus on the foot and ankle merely incorporate those structures into a grander appreciation of the body and insist that the body planes continue to run parallel to their torso definitions [26]. Conversely, some clinicians who focus on body imaging will make use of plane definitions that are more influenced by the imaging technology than by any definition of the anatomical position [27]. Foot and ankle specialists are often unconcerned with the conventional torso-based definitions, and apply the planes to the leg and foot in ways that seem to apply best to their research or clinical question. Some foot and ankle applications attempt to retain a single set of definitions for the anatomical planes by “turning the corner” at the ankle. When this is the case, subjects are essentially being forced into the HAP with a leg parallel foot. That solution, while often presented in a simplistic way, is actually the more complicated, and in some cases convoluted, approach. All of the objections to defining the HAP with a leg parallel foot would still apply. But, even more important, that solution seems to ignore the common ancestral/embryonic position. There is no fully satisfactory solution on how to define these planes. The error that can be avoided is the assumption that any one set of definitions will be universally understood.

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FIGURE 3.5 A comparison of the transverse and coronal planes as applied to the human condition projected onto the torso-based sagittal plane with the foot in the leg perpendicular (upper images) and leg parallel (lower images) orientations. When the foot is in the leg parallel orientation the transverse and coronal planes are roughly parallel to those as applied to the leg, thigh and torso. The definitions of these planes apply equally well to the foot, leg, thigh, and torso. The meanings of these planes are not as clear when applied to the foot in the leg perpendicular position. The plane that divides dorsal from ventral in the foot (coronal plane definition B) is roughly perpendicular to the leg, thigh and torso-based coronal planes. This suggests that the plane might be better described as the transverse plane of the foot. Similarly, the plane that divides proximal from distal in the foot (transverse plane definition 3) is at odds with its leg, thigh, and torso-based applications. As such, the concepts of the transverse and coronal planes can exchange meanings when applied to the foot in the leg perpendicular position. When motions are defined based on these planes, the definitions as applied to the foot can also change. As a result, there will be some circumstances where a single observed motion can be properly described by using two antithetical motion terms.

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3.2.5.3 Foot midline Most applications would expect the midline of the foot to run roughly parallel to the sagittal plane of the body, or the sagittal plane as defined in the LAP or FAP. Nonetheless, variation within these concepts will change the definition of the foot’s midline and its relationship with the sagittal plane (Fig. 3.6). The differing angular values produced by reference to these midlines are probably small, and in some cases may be smaller than intra-observer error. Therefore many will dismiss these differences as unimportant. In some applications that conclusion is probably correct, or at least defensible. But not for all, and therein lies the potential problem. An example of the problem associated with defining an appropriate midline can be found in the definition of the clinical condition “hallux valgus.” Most readers probably have a clear mental image of hallux valgus and are not confused by its presentation (Fig. 3.7). Confusion only occurs with the precise description of the pathology. Valgus is generally accepted to mean an outward bending away from the midline and varus is an inward bending. The problem with these descriptions is that the relevant midline is not clearly identified. Superficially, hallux valgus presents as a bulge at the first metatarsophalangeal joint, with a bend of the digit at this joint so that the hallucal phalanges angle toward the other digits. Less superficially obvious, but just as important to the condition, is the reformation of the first tarsometatarsal joint that angles the first metatarsal toward the body midline but away from the foot midline. This begs the question: to which bending joint surface does the condition hallux valgus actually refer? If we consider the hallux to be just the phalangeal part of the digit, then the bend at the metatarsophalangeal joint reorients the hallux away from the body midline (valgus) but toward the foot midline (varus). However, most digital positions are usually referenced to the midline of the foot (as in digital adduction-abduction). Continuity with this perspective would suggest that the pathological reorientation of the hallucal metatarsophalangeal joint is a varus condition. If we only consider the first tarsometatarsal joint, then the pathological hallucal metatarsal is body varus but foot valgus. There is generally not a lot of movement at the tarsometatarsal joints to provide a descriptive precedent here, however reference to the foot midline would seem to be more consistent. That would suggest that the pathological reorientation of the hallucal tarsometatarsal joint is the valgus condition, and therefore the change that is being described by the pathology’s name. Consultation with some available authorities does not resolve this issue. The American Orthopedic Foot & Ankle Society defines hallux valgus as “medial deviation of the first metatarsal and lateral deviation of the great toe” [28], which would seem to reference the HAP median plane. The American College of Foot and Ankle Surgeons [29] and the British Orthopeadic Foot & Ankle Society [30] both define the condition as an inward lean of the hallux. These descriptions, while admittedly for a nonspecialist audience, seems to confuse the concepts of varus and valgus, and contrast with recommended surgical correction focusing on the tarsometatarsal joint. The listing for “Hallux Valgus” in Medicine.net [23] also confuses the two midline concepts by defining hallux valgus as “a condition in which the big toe (hallux) is bent outward (toward the midline of the foot; valgus) so that it overlaps the second toe.” While none of these terminological confusions invalidate the recognition of this pathology, and would not alter standard radiographic measures of the deformity, they should raise an important question. If a simple condition, such as hallux valgus, is so difficult to describe with uniformity and precision, how can we be confident that we are understanding and explaining more complicated aspects of foot and ankle functional anatomy?

3.3

Foot motions

3.3.1 Defining motions A recent edition of Gray’s Anatomy [18] noted that “much confusion surrounds the descriptive terms for movement in the foot and ankle.” Much of this confusion can be linked to the variety of ways the cardinal planes are applied to the body, limb, and foot and the fact that most definitions of motion are linked to those planes. For the rotational motions: flexion-extension is generally described as occurring about an axis that emerges perpendicular to the sagittal plane; abduction-adduction is a rotation about an axis that emerges perpendicular to the coronal plane; and external-internal (or lateral-medial) rotation occurs either about an axis out of the transverse plane or about the longitudinal axis of the limb. Similarly, translations occur along these axes, or along an axis represented by the intersection of two cardinal planes. Because of the aforementioned ambiguities in the anatomical position, these motion labels can be problematic. However, even if the anatomical position problem is resolved, definitions of motions are still more complicated, especially when applied to the foot.

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FIGURE 3.6 Various applications of the sagittal plane concept, and hence the posterior anterior axis, as applied to the foot. Line A represents the sagittal plane of the body relative to a foot with forward oriented toes. Line B represents the sagittal plane of the body as it would appear relative to a foot with outward oriented toes. Both of these axis and plane definitions reflect conditions that would be employed during gait, or other whole body movement, studies where biomechanical considerations of the foot are considered as a component of a larger system. These two lines both represent the same, torso-based, sagittal plane. It is the changing orientation of the foot that makes them appear different. Lines C through E represent concepts of the sagittal plane as defined solely by reference to the foot (FAP), where each line represents a different concept for the midline of the foot. These three lines are all drawn to originate from the same point on the posterior calcaneus—although the actual location of this point is poorly defined and therefore probably varies considerably from study to study. Line C represents the midline definition that splits the second digit. The line through the second digital ray is often offered as the “functional” axis of the foot. In addition, the pedal dorsal interossei usually, but not always, reflect around the second digit. Line D passes through the gap between the second and third toes and represents the midline definition used by Inman [21]. Inman’s work is often cited in clinical literature, and therefore this midline definition could be thought of as the clinical standard. However, it is not clear that many clinicians actually use this midline definition. Line E represents the midline passing through the third digit. As such, it is the most obvious midline because it comes closest to splitting the foot into even right and left halves. Line F represents a posterior anterior axis as might be constructed in a leg-based anatomical position (LAP). In this case it represents one possibility for the floating posterior anterior axis as constructed by the ISB standard [20]. It runs from the midpoint between the two malleoli and is constructed to be perpendicular to the plantarflexion/dorsiflexion axis that is defined by the malleoli. While the given example projects Line F as being lateral to all the other constructions, because it is based on anatomical landmarks that are not found on the foot at all, its projection onto the foot would be a lot more variable across individuals than any of the other posterior/anterior axis constructs. Most applications that make use of this directional axis do not specifically reference the sagittal plane. However, motion definitions are applied in a way that treats this axis as being synonymous with the sagittal plane.

3.3.1.1 Flexion-extension In some texts, flexion is described as the motion that “bends a limb” or “decreases the angle” between segments [16,17,23]; extension would carry out the opposite movements. A slightly better definition, inasmuch as it would have better application in cross-species comparisons, makes reference to the original embryologic surfaces. Flexion is a rotation that attempts to bring a distal segment into closer proximity to the ventral surface of a more proximal segment; extension does the same with dorsal surfaces [18,31 33]. This concept is refined by referring to the limbs in their original developmental orientations, before embryological limb rotation. All embryonic limb depression motions are termed

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FIGURE 3.7 A foot with hallux valgus compared to two different concepts of the appropriate midline. This pathological condition is associated with an outward orientation at the tarsometatarsal joint and an inward orientation at the metatarsophalangeal joint. The pathology does not change based on the terminology used to describe it. Yet, it is unclear which joint reorientation is associated with the term “valgus”. Indeed, since both joint reorientations are common to this condition the name does not adequately describe the pathology.

joint flexion and all elevations are termed joint extension. This later developmental perspective has the added benefit of corresponding to neurological organization, so that all ventral division nerves innervate depressor muscles (all are flexor muscles in the foot and ankle system) and all dorsal division nerves innervate elevators (all are extensor muscles). This definition of movements is found in the primary actions of muscles such as flexor digitorum longus or extensor digitorum brevis. The problem with these definitions is that the embryological condition is not always well known to practitioners. Nonetheless, these definitions conform to most traditional descriptions of flexion and extension. However, most is not all. For example, Kapandji [34], a well-respected authority on joint physiology, defines flexion-extension of the foot and ankle as rotations in the opposite directions to how it is described here and would therefore have a muscle like flexor digitorum longus extend the toes. Definitions based on a biomechanical appreciation of the joints also seem to follow tradition more than a unifying logic. The grandest of traditions is that flexion-extension is a rotation that occurs about a roughly medial lateral oriented axis that emerges perpendicular to the sagittal plane. However, within most applications this flexion axis is usually defined within a kinematic reference frame and may therefore have only a casual relationship with the anatomical medial lateral axis. This occurs when the medial lateral axis is defined to be identical to the flexion-extension axis of the major joint in question. For example, in the ISB foot and ankle standard [20] the talocrural flexion-extension axis is treated as the definition of the medial lateral direction. This local reference frame medial lateral axis is rarely parallel to the medial lateral axis as defined within the torso. Many researchers are also aware that the orientation of a true rotational axis varies from individual to individual, and from joint to joint. However, when methods such as that suggested by the ISB standard are employed this variation is hidden. By definition the flexion-extension axis must always occur about the local reference frame’s medial lateral axis. Variation among individuals is now incorporated within the subject specific orientation of the local kinematic reference frame inside the anatomical position. Motion analysis studies often fail to account for this variation, because the concept of the anatomical position is supposed to be a standard, and therefore unvarying, orientation among individuals. The local kinematic reference frame is treated as an equivalent to the anatomical position, even though that is not the case.

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3.3.1.2 Adduction-abduction versus external-internal rotation Adduction is considered to be the rotational motion that occurs along the coronal plane and brings a limb segment closer to the midline; abduction takes the segment away from the midline. Traditional practice changes this concept when it is applied to movements of the toes. Even though the HAP may be in play, abduction-adduction of the toes, occurring at the metatarsophalangeal joints, brings the toes away from/toward the midline of the foot. This action occurs upon the plane of the metatarsals and thus about a ventral dorsal and inferior superior directed axis. The plane of the metatarsals is roughly equivalent to the HAP’s coronal plane when the foot is in a leg parallel alignment (Fig. 3.5). With the foot in a leg perpendicular orientation, this plane is more congruent to the HAP’s transverse plane. In the LAP or FAP this rotational axis can be seen as emerging from the coronal plane only when definition B, ventral dorsal separation, is applied. When abduction-adduction of the toes are defined as movements away-toward the midline of the foot, a conventional application of the right-hand rule for describing the rotation does not apply. According to the right-hand rule an abduction at the first metatarsophalangeal joint (rotating away from the midline of the foot) is a positive rotation, while an abduction at the fifth metatarsophalangeal joint is a negative rotation. It would seem that toe abduction-adduction is defined by tradition rather than by any logical concepts as applied to the rest of the body, although it does conform to traditional applications of abduction-adduction for the fingers. External-internal rotation is most often considered to occur about the long axis of the limb. The conventional application of this term does not typically reference the cardinal anatomical planes. Still, this concept is roughly similar to a rotation about an axis that emerges perpendicular to the transverse plane when that plane is defined to separate proximal from distal (definition 3). An important deviation from the cardinal plane concept is that along the leg these rotations are motions away from/toward the midline of the body as a whole. When applied to the foot these rotations can reference the midline of the body or they can reference the midline of the foot. Therein lies the complication and the confusion between external-internal rotation and abduction-adduction, and there can even be confusion about which specific joints are involved. Some researchers describe any out toe-in toe reorientation of the whole foot as abduction-adduction [34,35], even though the described motion may be occurring due to external-internal rotation at the hip or the knee. When this type of foot reorientation occurs at the talocrural joint, and perhaps in combination with the subtalar joint, terminology becomes situationally defined. Some practitioners continue to describe this motion as abduction-adduction, while others describe it as external-internal rotation (the perspective adopted for this book). In the perpendicular foot posture, this particular motion is not really possible as a primary result of muscular activity—there are no abductors-adductors of the human foot. However, when the human foot is oriented in the leg parallel posture, muscles that pull on the medial side of the foot (tibialis posterior and tibialis anterior) will create a foot adduction that is comparable to how the forearm muscles adduct the hand. The fibular (peroneal) muscles on the lateral side would display an abductor function. If the HAP is defined with a leg parallel foot then the terms abductionadduction could apply regardless of the orientation of the foot during data collection. Once again, the situation is dependent upon how the anatomical position is defined, and whether the implications of that definition are uniformly applied to all perspectives of the foot and ankle.

3.3.2 Whole foot motions and their complexity There are no single joint muscles that act upon the foot and ankle complex. While the isolated motions of individual joints can be studied and described, foot and ankle muscular activity does not affect isolated joints. Similarly, the ligaments that link bones at joints create a kinematic chain wherein the movement of one element necessitates a response in the other elements. Joint forces along the kinematic chain can be altered, or mitigated, by opposing muscular actions or by external reaction forces; but these opposing actions will have their own cascaded consequences. Added to this is the complexity associated with the HAP neutral foot positioned perpendicular to the leg. The result is a set of movement terminologies that are specific to the foot and ankle complex.

3.3.2.1 Plantarflexion-dorsiflexion Plantarflexion is generally considered to be the action that produces a simultaneous inferior displacement of the toes and superior displacement of the heel; dorsiflexion would be the opposite action [16]. Plantarflexion-dorsiflexion is often considered to be the primary action at the talocrural joint, although it is also not uncommon to see plantarflexiondorsiflexion used to describe other isolated joint actions throughout the foot [18]. When isolated joint actions are considered, plantarflexion-dorsiflexion can be seen as being directly synonymous with flexion-extension. However, because

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every muscular action creates simultaneous responses in multiple joints, a more complex foot encompassing terminology is required. The triceps surae muscle group are often considered to be the primary plantarflexors. These muscles insert onto the calcaneus and therefore have a direct and simultaneous influence on the talocrural and the subtalar joints. Due to the wedging of the trochlea talus, in combination with the tibiotalar and talofibular ligaments, this muscular action will also influence the tibiofibular joints. A cascade of distal trending ligaments (e.g., talonavicular, plantar calcaneocuboid, plantar calcaneonavicular, and bifurcate ligaments) will also cause this activity to influence the talonavicular and calcaneocuboid joints [36]. Ligaments trending distally from the navicular and cuboid will continue these, albeit diminishing, influences. The remaining muscles of the posterior leg compartment (tibialis posterior, flexor digitorum longus, and flexor hallucis longus) enter the foot by wrapping around the medial malleolus and therefore add a slight medially directed, inward, twist to the foot oblique to their flexing action on the joints they influence. Muscles of the lateral compartment (fibularis [peroneus] longus and brevis) wrap around the lateral malleolus and insert near the tarsometatarsal joints, and thereby contribute a slight flexion moment to several joints as they provide an outward twist to the foot. While these muscles would also influence the tibiofibular, talocrural, and subtalar joints, their more distal attachments will contribute more influence on the middle and distal foot joints. Jenkins [17] noted these relationships and defined plantarflexion as the combined influence of the posterior leg muscles upon the foot and ankle complex; dorsiflexion as the action of the muscles in the anterior leg compartment. When viewed in this way foot plantarflexion-dorsiflexion, while primarily a flexion-extension of the foot relative to the leg, would always be accompanied by an inward twist of the foot (Fig. 3.8). Therefore, the foot twist should be considered as a component of plantarflexion-dorsiflexion [37]. Because the fibular (peroneal) muscles are associated with an outward foot twist, their flexion-extension influences would need to be described by a different term. This muscular action based perspective contrasts with most kinematic assessments that focus on observed motions rather than identifying responses to specific muscular activity. Conventional kinematic data collection methods assesses ankle motion by evaluating rigid clusters placed upon the leg and the heel. This cluster arrangement encompasses two (talocrural and subtalar), if not more (tibiofibular), joints. It therefore captures the simultaneous actions of all of these joints. An a priori interpretation of kinematic data typically associates the talocrural joint with all the rotations observed along the sagittal plane and all other rotations are presumed to occur at the subtalar joint. Similar interpretations require that talocrural plantarflexion-dorsiflexion be accompanied by twists of the foot that occur in other, unspecified intrinsic, foot joints. These perspectives, by definition, restrict plantarflexion-dorsiflexion to the talocrural joint. The complexity of foot twisting that occurs along with talocrural plantarflexion-dorsiflexion are described by other terminologies.

3.3.2.2 Inversion—eversion One form of foot twisting is commonly associated with the terms inversion and eversion. However, what exactly is twisting and how that twisting is accomplished is not always clear. A survey of the literature [4] found that most

FIGURE 3.8 A comparison of foot orientations in plantarflexion (left image), neutral (central image) and dorsiflexion (right image). The plantarflexed foot displays an inward twist, slightly twisting the sole of the foot into the plane of the photograph. While the twist associated with the dorsiflexed position is not as obvious, it can be seen as a counter-twist from the plantarflexed position. Varying degrees of these twists accompanies most feet during plantarflexion-dorsiflexion. Therefore the definition of plantarflexion-dorsiflexion should not be limited to its actions along the sagittal plane (plane of the photograph). Instead, it might be better defined as a dimensionally complex action of the whole foot.

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published uses of these terms are not associated with a definition of their intended meaning. Of those surveyed publications that do include a definition, only 45% agreed that inversion-eversion is a rotation about the long axis of the foot so that the sole faces more medially-laterally. The second most common definition, found in 21% of the surveyed publications, describes a rotation about the subtalar joint so that the sole of the proximal (hind) foot faces more mediallylaterally. The important difference between these two definitions is that the first describes an action of the whole foot without attributing that action to any specific joint, while the second focuses on an action of the subtalar joint and its effect on the proximal (hind) foot. The remaining 34% of published definitions span 16 other meanings. Some published definitions are directly contradictory with others. While neither inversion nor eversion are included as “official” terms of motion within TA or NAV, the definition associated with a rotation about the foot’s longitudinal axis is found in many anatomy textbooks [16,18,33]. This definition avoids the terminological problems associated with requiring a definition of the foot’s coronal and transverse planes. Users need only establish a foot neutral reference plane, and for this it does not matter if the foot is in the leg parallel or perpendicular position. Applying the right hand rule defines inversion as a positive rotation of the foot plane and eversion as a negative rotation. When anatomical planes are defined, whole foot inversion-eversion is synonymous with either coronal plane adduction-abduction or transverse plane external-internal rotation. However, several questions arise with the longitudinal axis rotation definition. How is the rotation accomplished? What joints are involved? Is there only one way to accomplish the longitudinal rotation that is called inversion-eversion, or can this rotation be accomplished in several different ways by different combinations of intrinsic foot joint actions? If the latter, is it still correct to identify these differing combinations of joint actions using the same terminology? The varying answers to these questions can explain some of the diversity of inversion-eversion definitions that are found in the literature and can also explain why additional names for foot twists are frequently deemed to be necessary. Jenkins [17] described inversion as the combined action of the fibular (peroneal) muscles; eversion as the action of the peroneal muscles. The tibialis muscles primarily insert onto the first cuneiform, and so would initially act upon the first cuneonavicular joint, and secondarily upon the other joints crossed by their tendons (talonavicular, subtalar, and talocrural). As discussed earlier, the action of these muscle groups when acting upon the foot in a leg parallel position is roughly similar to the adduction actions seen in the wrist. But these are not the only actions reported for these two muscles. Tibialis anterior also contributes to dorsiflexion and tibialis posterior to plantarflexion, and it may be important to recall that some flexor muscles (flexor digitorum longus and flexor hallucis longus) also contribute to a medial twist of the foot. These related muscular actions, and anatomical plane definitions as applied to the foot, may be responsible for definitions of inversion-eversion that include mention of foot plantarflexion-dorsiflexion and abductionadduction [4,16]. Using a definition of inversion-eversion that separates movements of the proximal (hind) foot from movements of the distal (fore) foot suggests that each element is free to move independently. This would not discount instances where movement of the proximal foot could carry the distal foot with it; such a movement could still be described as an inversion-eversion of the proximal foot. Effectively this is a whole foot rotation about a longitudinal axis; an instance where the second most popular definition would incorporate the most popular definition. In other cases the foot could move so that the proximal and distal foot segments rotated in opposite directions relative to each other. When inversion-eversion is defined as a proximal foot rotation relative to the distal foot, an outward rotation of the distal foot could be seen as the same relative motion as an inward rotation of the proximal foot, even when the proximal foot is not actually moving. A further complication arises as to whether inversion-eversion should be used to describe motion about individual intrinsic joints other than the subtalar. The standard utilized in this book has these terms apply to individual joint rotations. But, what about joint motions that contribute to a whole foot inversion-eversion that cannot themselves be described as inversion-eversion motions? For example, sequential extension and/or flexion at the tarsometatarsal joints, MT1 to MT5, can raise the medial edge of the foot relative to the lateral edge and create the impression of the distal (fore) foot rotating about the longitudinal axis of the foot (distal foot inversion, but perhaps also proximal foot eversion). The opposite actions at these joints will create distal foot eversion. These actions correspond well with the definition of inversion-eversion as provided in an edition of Gray’s Anatomy [32]. It is important to reiterate that these are extension-flexion joint actions creating a whole foot motion that could be described as inversion-eversion. Indeed, rotation about a metatarsal longitudinal axis would not be forthcoming based on an assessment of most tarsometatarsal joint geometries. The question here is: are these tarsometatarsal joint motions properly described as inversion-eversion, a flexion-extension component of inversion-eversion, or as something else?

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3.3.2.3 Pronation—supination In anatomy textbooks pronation and supination are always described as an action of the radius upon the ulna that produces a longitudinal twist of the forearm. In many textbooks [16,17,25] the definition stops there, and in so doing imply that these terms do not describe actions in the foot. By analogy to the forearm, pronation-supination of the leg should describe movements between the tibia and fibula (replicating the movements at the radioulnar joints) or perhaps rotation about the longitudinal axis of the tibia on the femur (replicating the movement at the humeroradial joint) [4]. However, applications that make use of these definitions rarely appear in the literature. Instead, pronation-supination is used to describe a twisting action of the foot that has no analogy in the hand. Exactly what this twist is, and how or if it differs from the inversion-eversion twist, is extremely variable in its meaning, uses and definitions. A survey of the foot and ankle literature [4] found 21 competing descriptions of how these terms are applied. The most popular definition of pronation-supination, found in only 28% of surveyed articles, states that pronationsupination is: motion about the subtalar joint involving forefoot eversion-inversion, dorsiflexion-plantarflexion, and abduction-adduction. The subtalar joint is known to rotate about an axis that is oblique to all the anatomical planes [21,22]. This definition can be seen as an attempt to decompose the oblique subtalar rotation into three cardinal rotations. There are other published definitions that describe similar triplanar motion combinations without restricting motion to the subtalar joint or by using terms that differ perhaps because of their anatomical plane concepts. Assuming a triplanar decomposition, the most popular definition of inversion-eversion when interpreted as a whole foot motion seems to apply best; with the (subtalar) inversion-eversion component of pronation-supination applying to the overall inversion-eversion of the foot. The plantarflexion-dorsiflexion component is somewhat problematic, coming against the distinctions of describing whole foot versus joint specific motion. For the moment, substituting flexion-extension would resolve this particular problem. Finally, the typical confusions of abduction-adduction versus external-internal rotation can be avoided since however that component is labeled it would have to be orthogonal to the established eversioninversion and flexion-extension axes. The real problem with this pronation-supination concept, and definitions like it that are not limited to the subtalar joint, is that it remains an imprecise description of what is actually occurring. For example, when we encounter the frequent description “over pronation,” the phrase implies that the measurement of foot pronation exceeds some established standard. While the three rotations that make up this definition can be independently measured, the definition provides no concept of how they interact. Does the over pronation condition result only from an equal combination of excessive rotation in all three planes, or can it be caused by an over rotation in just one or two of those planes? Can an under rotation on one axis and an over rotation in another still qualify as over pronation? Is there some combination of over and under rotations about different axes that nullify each other? Has pronation-supination occurred when simultaneous rotations are observed to occur about only two axes, or even only one? The literature fails to answer these questions. As a result, these triplanar definitions of pronation-supination reflect impressions rather than useful measurable outcomes. When the literature reports measurements of foot pronation-supination very different definitions are used. The second most favored published definition of pronation-supination is only slightly less popular, being found in 25% of the surveyed reports. This definition is exactly the same as the most popular definition of eversion-inversion: rotation about the long axis of the foot so that the sole faces laterally-medially. This would lead to the impression that the two sets of terms are synonyms; which is an opinion shared by very few practitioners. Most of the published reports that use this definition of pronation-supination compare the relative twists (orientations) of the distal (fore) foot to the proximal (hind) foot. Sometimes these individual foot component twists are described by eversion-inversion, but the differences between them are described as pronation-supination. This can create circumstances where the proximal (hind) foot is described as undergoing pronation-supination solely due to a changing orientation of the distal foot that requires no motion in any component of the proximal foot. Presumably a similar situation could also occur with the distal (fore) foot pronating-supinating due to motion of the proximal (hind) foot. Therefore, when a foot is described as having pronated or supinated it can be unclear where and what motion has exactly occurred. All that can be certain is that the overall shape of the foot has changed. When one chooses to adopt these types of definitions, it should be noted that plantaris, soleus, and gastrocnemius are the only extrinsic foot muscles that attach to a bone (the calcaneus) of the proximal (hind) foot. Inasmuch as these muscles are predominately considered to be plantarflexors they are not suited to cause a twisting action of just the proximal (hind) foot. There are a few intrinsic foot muscles that attach to the calcaneus. While these muscles can produce twists of the calcaneus relative to the rest of the foot, they exhibit this capability under very limited and specialized

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circumstances. Their actions on the calcaneus would not be considered part of their normal function. In addition, no intrinsic or extrinsic foot muscles attach to the calcaneus or talus that might produce the opposing twist of the proximal (hind) foot. If pronation-supination is defined as motions of the proximal (hind) foot, it cannot be produced as a response to usual muscular activity. It then becomes an indirect motion that is produced as a by-product of more direct muscular actions. If such is the case, then pronation-supination becomes a constituent part of plantarflexiondorsiflexion, inversion-eversion or both. Kapandji [34] described this possibility when he defined inversion-eversion as a combination of adduction-abduction and supination-pronation.

3.4

Terminological implications of mathematical choices

Related to the concept of decomposing data into three individual rotations is the need to designate a rotation sequence. A rotation sequence must be specified to interpret, or decompose, the matrix into three simultaneous rotations about the mutually orthogonal X, Y, and Z axes of the reference frame. When kinematic data are gathered from rigid clusters, the changing relative orientations of those clusters are first encoded as a rotation matrix. The different choices for the orientation of the foot within the anatomical position, and the arbitrary nature of the related reference frame construction, will influence the expression of rotation values. Identical kinematic phenomenon will appear with different rotations when these reference positions are constructed differently (Table 3.1). However, even if anatomical orientation and reference frame variation are addressed, angular rotation values will also differ based on the presumed rotation sequence of the matrix. The rotation sequence specifies the order in which rotation values within a matrix are considered. Typically the X, Y and Z rotation values, known as Cardan angles or Tait Bryan angles (which are a subset of Euler angle rotations), are combined in one of six different sequences. The designation of the rotation sequence can be arbitrary and often varies among reports. The importance of the rotation sequence is generally well known among those in the biomechanics community, but often appears to be less appreciated by other consumers of kinematic data. The choice of rotation sequence can even seem inconsequential when the involved values are small—the rotation matrix associated with no motion will decompose into three zero value rotations regardless of the rotation sequence. However, as rotational values increase, different sequences will yield values that, although mathematically equivalent, can be widely different in their expressions (Table 3.2). Most kinematic reports mention the rotation sequence that was associated with its angular rotations. However, rotation values are often interpreted, compared, and applied without further reference to this sequence and without reference to the three linked rotations that were derived together. Because of this, it can be difficult to know that a 42.8 degrees rotation about the medial lateral axis from one report may be describing exactly the same phenomenon as a 17.2 degrees medial lateral axis rotation from another report (see XZY and YXZ sequence X axis rotation values in Table 3.2). When data are treated in their entirety, meaning proper simultaneous consideration of all three axis rotations and their associated sequence are made available, then it can be seen that the data are derived from the same rotation matrix and are therefore describing the same thing. This level of data reporting rarely occurs. Because of the apparent differences that can be associated with rotation sequences Woltring [38] proposed the attitude vector as a sequence independent method for reporting rotations. Despite its advantages the attitude vector has not become widely adopted and therefore effectively serves as a “seventh rotation sequence” that can be confused with the other six. Axis rotations are commonly treated in isolation, especially when we see plots or reports of values such as “average joint flexion.” Joint rotation values always exist in groups of three. For that reason, statistical summation must be derived from the simultaneous consideration of all three values—the rotation matrix [39,40]. Treating rotations as individual isolated values is mathematically invalid and therefore terminologically misleading. Regrettably this methodological error is so common, and unrecognized, that it has become the accepted method of kinematic data analysis and summation. Although we may eventually use identical terminologies that are based on matching definitions and anatomical perspectives, the absence of mathematically essential data will continue to muddle the interpretation of kinematic data.

3.5

Conclusion: standardizing foot and ankle terminology

The single most important barrier to the development of a standardized foot and ankle lexicon is the common belief that we already have one, the one we are using. It works. It always has worked. Why should it need to be modified or improved? Most of us recognize that different word choices and meanings occasionally crop up. But, when that

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TABLE 3.1 The effect of foot orientation within the anatomical position on the expression of rotations about the anatomical axes. Anatomical axis

Anterior foot orientation (degrees)

Lateral foot orientation (degrees)

X rotation (medial lateral axis)

15

2

Y rotation (posterior anterior axis)

18

23

Z rotation (inferior superior axis)

5

40

This table presents axis rotation values associated with the “average” subtalar joint function as described by Inman [22]. This average joint rotates 24 degrees about an axis that is oriented 42 degrees from the horizontal plane and 23 degrees from the foot midline. These data were used to create a rotation matrix that was then decomposed based on an XYZ sequence to obtain the rotation values for the anterior foot orientation. The rotation matrix was then rotated 33 degrees about the Z (inferior superior) axis to represent a laterally oriented foot. Rotation values were obtained based on an XYZ sequence decomposition from this reoriented matrix. Thus the reported rotation values represent the same joint motion that differ only by the orientation of the foot relative to the anatomical planes of the body. Notice that with the anterior foot orientation subtalar joint motion would primarily be associated with rotations around the medial lateral and posterior anterior axes. Viewing the foot in a laterally oriented, out-toed, posture changes the interpretation of subtalar joint motion so that it now appears to occur primarily about the posterior anterior and inferior superior axes. In both cases, the subtalar joint motion was exactly the same. It is only the difference in the foot’s orientation within the anatomical reference frame that creates the dramatic differences in the reported rotation values.

TABLE 3.2 Comparison of individual axis rotation angles derived from the same rotation matrix. Each set of angles differ only by the rotation sequence used to decompose the matrix or by the use of the attitude vector expression. 8 < 0:66; Rotation matrix 0:74; : 2 0:14;

2 0:56; 0:61; 0:56;

9 0:50 = 2 0:30 ; 0:81

Rotation sequence

Rotation about X axis (degrees)

Rotation about Y axis (degrees)

Rotation about Z axis (degrees)

XYZ

20.0

30.0

40.0

XZY

42.8

37.0

33.8

YXZ

17.2

31.6

50.3

YZX

25.9

11.9

47.3

ZXY

34.3

9.8

42.4

ZYX

34.7

8.1

47.9

Attitude vector

29.2

21.8

43.9

happens, it is the other person who is incorrect, or at best playing loose with accepted definitions. We reinforce this attitude by working in groups that share vocabulary. When we repeat our words, and their meanings, enough times within our group we convince ourselves that we are indeed correct. Repeatedly using a term in a specific way by many people over a long period of time may eventually earn acceptance for that usage within the community. Until then, the usage is generally considered to be incorrect. This is how words and their meanings are added to any language. The problem here is that there are several parallel communities that contribute to and consume the foot and ankle literature. Major findings are almost certainly communicated across the communities, but the day to day accumulation of smaller research and aspects of clinical practice are not. This allows for each community, and sometimes each laboratory, clinic or university, to develop its own unique terminological definitions. No group is “wrong” in how it defines its terms. However, the group membership is wrong when they assume that their definitions are the only correct definitions and that every other group will understand terms in exactly the same way.

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The final stage of the scientific process is the successful communication of results and conclusions. The diversity of meanings and perspectives associated with the terminology of specialized disciplines makes clear communication difficult. Several authorities have attempted to solve this problem by creating lists of terminological standards. It is unfortunate that many of these attempts stand at cross purposes. The net result is that existing terminological standards do not bridge the gaps between different intellectual perspectives, but instead seem to reinforce them. We can ask the following: can there be a standardized foot and ankle terminology? Given the terminological diversity that has been outlined in this chapter, and the ideological momentum that supports it, the answer to that question appears to be “no”. That may be the wrong question. Perhaps instead we should ask: should there be a standardized foot and ankle terminology? The answer to that question is probably also no. But that does not mean that we must condemn ourselves to a morass of foot and ankle miscommunication. What we need to do is to agree that since terminological standards do not exist we cannot assume that technical terms will be obviously understood in the same way by all people. The solution may be that we agree to define all of our technical terms as we use them. Defining everything in every publication would rapidly become wearisome. Eventually, readers would gloss over the definitions and reviewers and editors may not give them close attention. However, journal and published volumes, especially those that focus on specialized topics such as the foot and ankle, could alleviate this problem by publishing a terminological glossary in their front matter at least once each volume, if not in every issue. Publishing this glossary would be akin to the current author guidelines that limit submissions to a particular language. Authors, editors, reviewers, and readers would be encouraged, or rather required, to reference this list as the journal’s lingua franca. For a foot and ankle publication this list might include items like: G G G

G G G

G

Acceptable eponyms, and perhaps words or concepts that should not be associated with eponyms. Names for limb segments, such as Leg versus Shank or Hallux versus First Digit. Description of how the anatomical position is defined for the foot and for the foot and leg, and how those definitions relate to the body as a whole. How the anatomical planes are defined for the foot and leg and how they relate to the body as a whole. How anatomical directions are to be defined relative to the foot and to the body. How movements will be defined according to these definitions and if some movement terms are to be considered joint specific or apply to the foot as a whole. A rotation sequence that will be used to describe and analyze joints. Either a universal sequence or a specific sequence for each joint.

Adoption of a recommendation like this will probably not be imminent. When we act as authors, word count limits and reviewer tolerance will prevent us from defining every specialized term in our contributions. In fact, we may even learn to play the publication game by not defining a specialized term and allow the editor and reviewers to assume that their definitions are the same as ours. So, we must guard ourselves against this behavior when we are acting as readers and consumers of research reports. We need to question, and when possible set aside, our own definitions and try to glean meanings from the context clues left by the authors. The only guard we have against the problems of a nonstandardized terminology is the ability to understand where, and how, misunderstandings can occur.

References [1] Federative Committee on Anatomical Terminology. Terminologia Anatomica. Thieme, Stuttgart; 1998. [2] International Committee on Veterinary Gross Anatomical Nomenclature. Nomina Anatomica Veterinaria. Hannover: World Association of Veterinary Anatomists; 2012. [3] Huson A. Joints and movements of the foot: terminology and concepts. Acta Morphol. Neerl.-Scand. 1987;25:117 30. [4] Greiner TM. The jargon of pedal movements. Foot Ankle Int 2007;28:109 25. [5] Kachlik D, Bozdechova I, Cech P, Musil V, Baca V. Mistakes in the usage of anatomical terminology in clinical practice. Biomed Pap Med Faculty Univ Palacky´, Olomouc, Czechoslovakia 2009;143:157 62. [6] Martin BD, Thorpe D, Merenda V, Finch B, Anderson-Smith W, Consiglio-Lahti Z. Contrast in usage of FCAT-approved anatomical terminology between members of two anatomy associations in North America. Anat Sci Educ 2010;3:23 32. [7] Woywodt A. Should eponyms be abandoned? Yes. Br Med J 2007;335:424. [8] Fargen HM, Hoh BL. The debate over eponyms. Clin Anat 2014;27:1137 40. [9] Olry R. Anatomical eponyms, part 2: The other side of the coin. Clin Anat 2014;24:1145 8. [10] Whitworth JA. Should eponyms be abandoned? No. Br Med J 2007;335:425. [11] Olry R. Anatomical eponyms, part 1: to look on the bright side. Clin Anat 2014;24:1142 4.

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[12] Thomas PBM. Are medical eponyms really dying out? A study of their usage in the historical biomedical literature. J R Coll Phys Edinb 2016;46:295 9. [13] Dictionary.com. https://www.dictionary.com. [accessed 23.04.18]. [14] Gray H. Anatomy: descriptive and surgical. London: John Parker & Son; 1858. [15] Cunningham DJ. Manual of practical anatomy. London: Young J. Pentland; 1893. [16] Gray H. In: Clemente C, editor. Anatomy of the human body. 30th ed. Baltimore, MD: Williams & Wilkins; 1985. [17] Jenkins DB. Hollinshead’s functional anatomy of the limbs and back. 9th ed. Philadelphia, PA: Saunders; 2009. [18] Standring S, editor. Gray’s anatomy: the anatomical basis of clinical practice. 39th ed. Edinburgh: Elsevier; 2005. [19] Moore KL, Dalley AF, Agur AMR. Clinically oriented anatomy. 7th ed. Baltimore, MD: Lippincott Williams & Wilkins; 2014. [20] Wu G, Siegler S, Allard P, Kirtley C, Leardini A, Rosenbaum D, et al. ISB recommendation on definitions of joint coordinates system of various joints for the reporting of human joint motion—part I: ankle, hip and spine. J Biomech 2002;35:543 8. [21] Isman RE, Inman VT. Anthropometric studies of the human foot and ankle. Bull. Prosthet. Res. Spring; 1969, p. 97 129. [22] Inman VT. The joints of the ankle. Baltimore, MD: Williams & Wilkins; 1976. [23] Medicine.net, http://www.medicinenet.com [accessed 15.03.18]. [24] Drake RL, Vogl AW, Mitchell AWM. Gray’s basic anatomy. Philadelphia, PA: Elsevier; 2012. [25] Moore KL, Agur AMR, Dalley AF. Essential clinical anatomy. 5th ed. Philadelphia, PA: Wolters Kluwer; 2015. [26] El-Khoury GY, Bergman RA, Montgomery WJ. Sectional anatomy by MRI/CT. New York: Churchill Livingstone; 1990. [27] Rubin DA, Towers ID, Britton CA. MR imaging of the foot: utility of complex oblique imaging planes. Am J Radiol 1996;166:1079 84. [28] American Orthopaedic Foot & Ankle Society. Hallux Valgus, ,http://www.aofas.org/PRC/conditions/Pages/Conditions/Hallux-Valgus.aspx. [accessed 15.03.18]. [29] American College of Foot and Ankle Surgeons. Bunions (Hallux Valgus), ,https://www.acfas.org/footankleinfo/bunions.html. [accessed 15.03.18]. [30] British Orthopeadic Foot & Ankle Society. Hallux Valgus (Bunion), ,https://www.bofas.org.uk/Patient-Information/Hallux-valgus-bunion. [accessed 15.03.18]. [31] Romanes GJ. Cunningham’s manual of practical anatomy. 15th ed. Oxford: Oxford University Press; 1986. [32] Williams PL, Warwick R, Dyson M, Bannister L. Gray’s anatomy. 37th ed. Edinburgh: Churchill Livingstone; 1989. [33] Rosse C, Gaddum-Rosse P. Hollinshead’s textbook of anatomy. 5th ed. Philadelphia, PA: Lippincott-Raven; 1997. [34] Kapandji IA. (trans.) In: Honore LH, editor. The physiology of the joints. Vol 2 Lower limb. 6th ed. Edinburgh: Elsevier; 2011. [35] Sammarco GJ, Hockenbury RT. Biomechanics of the foot and ankle. In: Nordin M, Frankel VH, editors. Basic biomechanics of the musculoskeletal system. 3rd ed. Philadelphia, PA: Lippincott Williams & Wilkins; 2001. [36] Lundberg A, Goldie I, Kalin B, Selvik G. Kinematics of the ankle/foot complex: plantarflexion and dorsiflexion. Foot Ankle 1989;9:194 200. [37] Greiner TM. Correlated joint rotations in the medial foot and the definition of plantarflexion-dorsiflexion. J Foot Ankle Res 2012;5:O41. [38] Woltring HJ. Representation and calculation of 3-D joint movement. Hum Mov Sci 1991;10:603 16. [39] Pierrynowski MR, Ball KA. Oppugning the assumptions of spatial averaging of segment and joint orientation. J Biomech 2009;42:375 8. [40] Greiner TM, Ball KA. Statistical analysis of the three dimensional joint complex. Comput Methods Biomech Biomed Eng 2009;12:185 95.

Chapter 4

Kinematics and Kinetics of the Foot and Ankle during Gait Jason T. Long1,2 and Joseph J. Krzak3,4 1

Motion Analysis Lab, Cincinnati Children’s Hospital Medical Center, Cincinnati, OH, United States, 2Department of Orthopaedic Surgery,

University of Cincinnati, College of Medicine, Cincinnati, OH, United States, 3Physical Therapy Program, Midwestern University, Downers Grove, IL, United States, 4Motion Analysis Center, Shriners Children’s, Chicago, IL, United States

Abstract During gait, the foot acts as the primary point of contact between the body and the ground. Movement and loading patterns of the foot and ankle joints directly impact the quality of a person’s motion and the efficiency of ambulation. To understand and effectively treat any level of ambulatory impairment, it is therefore necessary to have a thorough understanding of foot kinematics and kinetics. This chapter explores the articulatory anatomy of the foot and ankle, and typical patterns of motion and moments during gait. It also highlights different approaches for measuring these values, as well as emerging opportunities to gain greater insight into the workings of this very complex mechanism.

4.1

Introduction

According to Perry and Burnfield, each weight-bearing limb accomplishes four distinct functions during locomotion: (1) upright stability is maintained despite an ever-changing posture, (2) progression is generated by the interaction of selective postures, muscle forces, and soft tissue elasticity, (3) floor impact at the onset of each stride is minimized, and (4) energy is conserved by performing these functions in an efficient manner that minimizes the amount of muscular effort required [1]. The simultaneous accomplishment of these four functions depends on distinct motion patterns which represent a complex series of interactions between the upper and lower body and the ground. The foot and ankle play an integral role in these functions, acting as the principal interface between the individual and the support surface, as well as a mechanical link to more proximal segments. Indeed, in a span of slightly over a second, the foot of a healthy ambulator experiences the impact of initial foot strike, the weight of the body during first dual limb and then single limb stance, the push off that marks the end of stance phase, and the brief period of rapid limb advancement during swing phase. Given its breadth of roles in load accommodation, stability, and progression, over the past decades clinicians and researchers have subjected the foot to increasing levels of clinical and experimental scrutiny. While in contact with the support surface, the foot is required to perform several intricate functions allowing for both stability and mobility while minimizing energy expenditure. Stability throughout the joint complexes of the foot and ankle requires the functional integration of proprioceptive feedback, joint mobility, and muscle control. During stance phase, motion at the foot and ankle propagates rotation of more proximal limb segments. This motion also minimizes force generation requirements from more proximal muscles by providing a stable base during forward progression [2]. From a temporal perspective, the gait cycle is typically broken up into seven biomechanically relevant phases. At the foot and ankle, however, stance phase is also described in terms of three “rockers” of gait. These rockers describe fundamental components of sagittal plane motion during the gait cycle (Fig. 4.1), and deficits in motion or timing during these rockers are generally linked to pathology or ambulatory dysfunction. From initial contact through midstance (commonly referred to as the first and second rocker, respectively), flexibility of the structures within the foot allow it Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00020-2 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 4.1 Normal sagittal plane ankle kinematics associated with a lumped ankle joint and single segment foot model (foot relative to tibia, calculated via a sagittal-coronal-transverse Euler sequence). Data are presented in a style typical of clinical gait analysis. The shaded band represents data (mean 6 1 SD) from a population of healthy ambulators (n 5 52 M, 56 F; age 10.8 6 4.0 years). The single lines represent bilateral information from a representative single healthy ambulator (11 years female). The red (solid) line indicates data collected from the right ankle, and the blue (dashed) line indicates data from the left ankle. Positive values indicate dorsiflexion. Normalized time (% stride) is represented on the x-axis. Stance phase is further subdivided into first rocker (yellow), second rocker (green), and third rocker (blue), with associated foot positions depicted.

to absorb forces from the support surface and accommodate to alterations in terrain [2]. From terminal stance through the transition into swing phase (third rocker), stability of the structures within the foot facilitates transformation of the foot into a rigid structure that allows for power generation at the ankle to accelerate the body forward (Fig. 4.1).

4.2

Overview of relevant anatomy

Two individual joints make up the articulation commonly referred to as the “ankle joint complex.” The more proximal of these is the talocrural (also tibiotalar) joint [3]. The trochlea of the talus fits into the mortise formed by the distal tibia/fibula, with loadbearing taking place between the tibia and the domelike superior surface of the talus. With no muscles directly inserting onto the bone, the talus is constrained by the medial and lateral malleolus and various ligaments and tendons. As such, the joint operates predominantly as a hinge joint. Other aspects of talar geometry provide additional elements of stability; for example, the talar width gradually increases in the anterior direction, such that maximum stability is achieved in a dorsiflexed position [4]. These aspects of bone geometry and articular configuration lead to the talus’ primary role in plantarflexion and dorsiflexion of the foot during gait. However, the axis of rotation of the ankle joint is externally rotated at approximately 19 degrees (range 15 25) in the transverse plane from pure sagittal. As a result, the talocrural joint likely does not operate as a simple hinge joint, and some motion is available outside of the sagittal plane [5,6]. Motion at the subtalar joint significantly contributes to achieving the goals of both stability and mobility of the foot during locomotion. However, because of the talus’ unique architecture, the location of the subtalar joint’s oblique axis of rotation (including antero-posterior, vertical, and transverse components), inter-subject variability of the location of the subtalar axis, and the inability to track the talus using traditional motion capture techniques, calculating motion of the subtalar joint during locomotion is technically difficult. Further, the talus lacks bony landmarks that can be reliably palpated, so approximating subtalar motion using external (surface) markers is usually based on various rigid body modeling assumptions. Subtalar kinematics are often combined with talocrural motion to produce hindfoot kinematics. Much of what is currently known about the specific kinematics of the subtalar joint during locomotion comes from cadaveric and intracortical bone pin studies [7,8] based on previous work by Inman and Mann [9,10]. Motion about the subtalar joint has been compared to a mitered hinge [9,10]. During first rocker and into the early portion of second rocker (approximately 10% of stance phase), the leg internally rotates. Internal rotation of the leg results in calcaneal eversion with plantarflexion and internal rotation of the talus, producing a supple foot that is capable of absorbing shock during initial contact and accommodating to support surface irregularities during early stance phase. From second rocker through third rocker, the leg externally rotates. External rotation of the leg results in calcaneal inversion with dorsiflexion

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and external rotation of the talus, and converts the foot to a rigid lever capable of transmitting forces from the support surface, propagating the center of mass over a stable base of support, and generating power to accelerate the body forward. Although separate articulations, the calcaneocuboid and talonavicular joints are often combined and referred to as the transverse tarsal joint due to their anatomical and functional relationships. Motion at the transverse tarsal joint is typically coupled to motion at the more proximal joints, specifically the subtalar joint. During ambulation on a level surface, multiplanar motions of the transverse tarsal joint are typically opposite of motions at the subtalar joint, with an ultimate goal of achieving a plantigrade foot with a balanced force distribution throughout the foot in stance phase [10]. From first rocker into the early portion of second rocker, the leg rotates internally, the calcaneus everts, and the talus plantarflexes and rotates internally. This proximal motion results in a parallel orientation of the axes of the calcaneocuboid and talonavicular joints. When these axes are parallel, the midfoot unlocks and additional motion becomes available through the transverse tarsal joint, further contributing to the foot’s suppleness and its ability to absorb shock and conform to irregularities in the support surface. During this portion of stance phase, the transverse tarsal joint goes through a combined inversion/flexion/internal rotation relative to the talus and calcaneus. This reciprocal motion (relative to the hindfoot) creates a plantigrade foot, with the fifth metatarsal head in contact with the floor [11]. From second rocker through third rocker, the leg externally rotates, the calcaneus progresses into inversion, and the talus dorsiflexes and externally rotates [11]. During this phase, the calcaneocuboid and talonavicular axes are no longer parallel. Motion about the transverse tarsal joint is limited by this orientation, which contributes to the foot’s rigidity and ability to generate power at the ankle. To maintain the first metatarsal head’s position on the ground and balanced weight distribution throughout the foot, the transverse tarsal joint initially everts, extends, and externally rotates relative to the hindfoot. However, this transverse tarsal joint motion is minimal due to the orientation of the calcaneocuboid and talonavicular axes. As the leg, the calcaneus, and the talus progress into greater inversion, extension, and external rotation, the transverse tarsal joint moves along with the more proximal segments (inversion, flexion, and internal rotation). In the event of either extreme coronal hindfoot position or limitations in transverse tarsal joint motion, the tarsometatarsal joints (primarily the first and fifth) account for the remaining sagittal and coronal motion needed to achieve a plantigrade foot during stance phase [12]. The first metatarsophalangeal (MTP) joint is capable of both abduction/adduction and flexion/extension, but the majority of its motion during locomotion occurs in the sagittal plane. In later stance phase (late second rocker through third rocker), the first MTP joint extends as the heel lifts off of the ground. Extension of the first MTP joint results in tensioning of the plantar fascia referred to as the “windlass effect” [13]. This tensioning further stabilizes the foot so that it can contribute to power generation at the ankle. While most feet will follow these general patterns, newer evidence from biplane fluoroscopy and bone pin studies have found that significant variability in kinematics exists even between feet that are considered to be healthy and neutrally aligned.

4.3

Overview of kinematic and kinetic modeling

This functional understanding of foot motion and loading can be enhanced through the use of quantitative three-dimensional (3D) motion analysis. In the clinical context, 3DMA typically involves the simultaneous collection of temporal-spatial, kinematic, kinetic, and electromyographic data during locomotion. Quantitative gait analysis has been extensively utilized to study individuals across a wide range of ages and pathologies, helping to guide clinicians’ decisions about appropriate interventions, providing objective data about the efficacy of orthopedic surgery and/or conservative treatments, and longitudinally tracking the progression of various disease processes [14]. A common method for measuring kinematics during quantitative gait analysis involves the use of motion capture cameras to track small reflective markers affixed to anatomically relevant points on a limb segment of interest. A minimum of three non-collinear markers are required to track the position and orientation of any rigid body (i.e., limb segment). The orientation of the rigid body in question can then be calculated relative to another limb segment or to the global coordinate space. Traditionally, kinematics are reported in degrees of motion, with time normalized to 100% of the gait cycle (the timespan from initial foot contact to the next ipsilateral initial contact, e.g., Fig. 4.1). Traditional 3D motion analysis often tracks the pelvis, thigh, and shank segments, as well as the combined head, arms, and torso [15]. Early work in this area (as well as current work that does not focus on foot kinematics) often uses a single rigid body to represent the foot, with a revolute joint between the tibia and foot representing the ankle [15]. While not anatomically correct, this single-segment model of the foot does provide a simplified and relatively easy way to interpret perspective on foot and ankle mechanics. Despite the limitations, a wide range of published studies have used this simplified approach for characterization and follow-up studies involving orthopedic surgery [16 20], neurosurgery [21], and physiatry [22], as well as investigations of orthotic intervention [23,24] and prosthetic design [25,26]. Indeed, as long as its limitations are documented and understood, a model that uses a single-segment representation of the foot can provide a great deal of useful data, particularly when the research questions are more proximal in nature.

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One approach in marker-based motion capture would involve surface markers placed on the lateral shaft of the tibia, lateral malleolus, and the midpoint of the second and third metatarsals. Prior to walking trials, a static calibration trial is collected to create additional “virtual” markers located within the knee and ankle joint centers. These markers are created using the location of the known surface markers along with subject-specific anthropometric measurements and function to identify the segment in 3D space, mathematically linking the surface markers with the underlying skeletal anatomy [27]. The motion of the foot during walking (dynamic) trials is based on a vector that passes from the calculated ankle (talocrural) joint center through the marker placed over the second and third metatarsal heads. The plantar surface of the foot provides a reference for the sagittal plane alignment of this vector, assuming a plantigrade foot. The ankle center and toe marker are intended to provide references for transverse plane alignment. The foot and ankle kinematics calculated using the method just described are two-dimensional, providing measures of sagittal and transverse plane motion of the foot relative to the tibia. An additional measure of foot progression angle is also calculated to describe the transverse plane orientation of the foot in relation to the global coordinate axes [15]. This process of measuring motion, more specifically rotations, of a distal segment (foot) in relation to the more proximal segment (tibia) is performed through a series of limb rotation algorithms that approximate clinically relevant anatomical planes of motion [28]. While the transverse plane position of the foot is relatively static throughout the stride, the sagittal plane motion demonstrates the three distinct rockers that were noted earlier. During the first rocker, the ankle plantar flexes as the foot contacts the ground under the control of the anterior muscles. The posterior location of the ground reaction force (GRF) and the weight of the foot result in an internal dorsiflexion moment. As the foot and ankle progress into the second rocker, the ankle plantarflexors control the forward progression of the tibia over the foot resulting in increasing dorsiflexion. As the body advances over the foot, the location of the GRF travels anteriorly through the ankle joint center to the forefoot and hallux resulting in a steady increase in an internal plantarflexor demand (Fig. 4.2). Power is absorbed through second rocker, resulting in a controlled forward progression as the limb/body decelerates in the anterior direction. The third rocker is characterized by rapid plantarflexion of the ankle that provides 60%-80% of the propulsive power generated by the lower limb, energy which is used to accelerate the limb/body in the anterior direction. During the swing phase of gait, activation of the anterior muscles results in ankle dorsiflexion to provide adequate swing phase clearance of the toes, and to appropriately preposition the foot prior to another cycle of gait. The overall excursion of ankle motion in the sagittal plane is approximately 30 degrees. Transverse plane foot kinematics comprises both ankle rotation (foot relative to the tibia; Fig. 4.3A) and foot progression angle (foot relative to the forward line of progression created by the global coordinate axes; Fig. 4.3B). Minimal motion (approximately 6 degrees) is observed at the foot in the transverse plane. The foot progression angle (i.e., the transverse plane position/motion of the foot relative to the global coordinate axes) is between 5 and 20 degrees of external rotation. Foot rotation (i.e., the transverse plane position/motion of the foot relative to the bimalleolar axis) is generally neutral. The foot progresses into slight external rotation during first and early second rocker as it maintains its flexibility so as to accommodate to the support surface. From second rocker through third rocker, the foot progresses toward internal rotation as it transitions into a rigid lever capable of generating power to accelerate the limb/body in the anterior direction. FIGURE 4.2 Normal sagittal plane ankle kinetics associated with a lumped ankle joint and single segment foot model (foot relative to tibia, calculated via a sagittal-coronal-transverse Euler sequence). Data are presented in a style typical of clinical gait analysis. The shaded band represents data (mean 6 1 SD) from a population of healthy ambulators (n 5 52 M, 56 F; age 10.8 6 4.0 years). The single lines represent bilateral information from a representative single healthy ambulator (11 years female). The red (solid) line indicates data collected from the right ankle, and the blue (dashed) line indicates data from the left ankle. Positive values indicate internal plantarflexion demand. Normalized time (% stride) is represented on the x-axis. Stance phase is further subdivided into first rocker (yellow), second rocker (green) and third rocker (blue).

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FIGURE 4.3 Normal transverse plane ankle kinematics (A: foot relative to tibia) and foot progression angle (B: foot relative to line of progression) associated with a lumped ankle joint and single segment foot model. Data are presented in a style typical of clinical gait analysis. The shaded band represents data (mean 6 1 SD) from a population of healthy ambulators (n 5 52 M, 56 F; age 10.8 6 4.0 years). The single lines represent bilateral information from a representative single healthy ambulator (11 years female). Positive values indicate internal rotation. Normalized time (% stride) is represented on the x-axis.

4.4

Healthy and impaired feet

It is commonly assumed that deviations in kinematic, kinetic, and EMG data (when compared to a healthy population) result in a less efficient gait. Ballaz et al. examined the relationship between gait efficiency and lower limb kinematics during walking [29]. Assessing adolescents with cerebral palsy (CP), the authors compared a calculated energy expenditure index, based on the linear relationship between heart rate and oxygen consumption, to peaks of sagittal plane joint kinematics during locomotion. Strong correlations existed between energy expenditure index, ankle range of motion (r 5 20.70, P , .02) and peak ankle plantarflexion (r 5 0.74, P , .05) while walking. These findings suggest that ankle motion during gait is a key kinematic factor associated with gait efficiency among adolescents with CP. Brodsky et al. studied stage II posterior tibial tendon dysfunction (PTTD) in a series of 12 patients who underwent tendon substitution and medial displacement calcaneal osteotomy [30]. They used a five-camera motion analysis system to evaluate foot and ankle kinematics pre- and post-operatively, using a single segment to represent the foot. Despite finding reductions in cadence, step length, velocity, and ankle push-off power, they found that the overly simplified foot model was insufficient to study the characteristics symptoms of PTTD (excessive hindfoot valgus and forefoot external rotation). A slightly more complex version of this approach was adopted by Thomas et al. in a study of patients following ankle arthrodesis [31]. The authors placed markers at the shaft and heads of the first and fifth metatarsals to represent the forefoot, and modeled this segment as neutrally aligned during comfortable stance. This study was similarly hampered by an inadequate level of detail in foot motion, and though they found reduced ranges of motion in the arthrodesis patients compared to a group of controls, they could not discern any positional differences between the groups. While a single segment representation provides the simplest and most clinically straightforward means of tracking foot motion, the limitations of such an approach are clear. This model leads to relative motion between the tibia and foot segments being represented as “ankle motion” (grouping the talocrural and subtalar joints into a single “ankle” joint), and motions between the intrinsic bones of the foot are disregarded altogether. When the multiple articulations between these intrinsic bones are not taken into account, the transition of the foot from a supple load attenuator (during initial contact and loading response) into a rigid lever (during preswing and at toe-off) cannot be appreciated. In patients with foot pathology or deformity, abnormal positions and/or motions of the distal tarsals and metatarsals can corrupt calculations of ankle motion, and the very motions which characterize the pathology can be masked by the generalizations built into the model. Davis et al. described one instance of this in patients with flatfoot, in which findings of excessive dorsiflexion did not reflect actual ankle motion, but instead stemmed from flattening of the longitudinal arch [32].

4.5

Multisegment foot models

To fully appreciate the role of the foot and ankle during locomotion, a more detailed model is necessary (Fig. 4.4). The concurrent goals of providing both stability and mobility at the foot during locomotion are accomplished by the complex skeletal architecture comprised of 26 bones, 33 articulations, several dozen muscles, and scores of ligaments [33].

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FIGURE 4.4 Schematic of single segment foot model (left foot) versus multisegmental foot model (right foot). The Milwaukee Foot Model is used as an example for this schematic, with segments including hindfoot, forefoot, and hallux (tracked via triad affixed to the proximal phalanx, which creates the “floating” segment in the image). Tibia segment on the left side has been omitted from this image for ease of visualization.

To simplify the complexity of the foot, the bones and their associated articulations are commonly grouped into clinically relevant segments: G G

G

G

The hindfoot segment is comprised of the talus and calcaneus and includes the talocrural (ankle) and subtalar joints. The midfoot segment is comprised of the navicular, cuboid, and the cuneiforms. The midfoot has proximal articulations with the talus (talonavicular joint) and the calcaneus (calcaneocuboid joint), as well as interarticulations among the bones. The forefoot segment includes the metatarsals and phalanges. The forefoot has proximal articulations with the cuboid and the cuneiforms. As a result of pathology and interventions unique to the first phalanx, the hallux is also commonly modeled as an individual segment.

The challenges involved in dividing the foot in this manner are multifaceted. They include the small size of many of the bones of the foot, which can make affixing multiple markers difficult or impossible. A similar problem is encountered with the talus, which plays a critical role in the ankle joint complex (the talocrural and subtalar joints), but which cannot be palpated or marked outside of a position of excessive plantarflexion. The oblique nature of some of these intersegmental articulations, such as that between the talus and calcaneus (i.e., the subtalar joint) can also complicate the creation of a model that is clinically meaningful. Additionally, the rigid body assumptions that are typical in lower extremity modeling come into question when applied to the multiple segments of the foot and ankle. This is especially notable in the forefoot, where multiple bony segments and intertarsal articulations introduce flexibility that violates the rigid body model assumptions. On top of the modeling complexities involving bones, there are an array of soft tissue challenges in the foot, where over 30 muscles and 100 ligaments take part in the cyclic loading and unloading activities of gait. Difficulties can arise in tracking when a muscle such as extensor digitorum longus contracts, and the resulting tension changes the position of a marker or sensor placed on the distal portion of the foot. A variety of multisegmental approaches to foot and ankle modeling have been published and described in detail elsewhere in this book. Fundamental to the definition of such models, however, is the concept of the neutral alignment or “zero” pose/position. In most clinically relevant models, when the orientation of a distal segment is measured relative to a more proximal segment, neutral alignment of the segments is indicated by a measure of zero in all three

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planes. This is conceptually simple at the more proximal joints, as the single-segment nature of pelvis/thigh and thigh/shank alignment can be easily visualized. When the foot is visualized as a single rigid segment, this conceptual simplicity persists at the ankle (with the understanding that a foot that is “aligned” with the leg is actually positioned at a right angle to it). When the foot is subdivided into multiple segments, however, a more nuanced definition of zero position becomes necessary. This is partly due to the fact that, unlike in the thigh and shank, the orientations of the tarsal segments are not clearly apparent from visual inspection of the foot. In the absence of easily palpable landmarks, most modeling solutions for the multisegmental foot adhere to one of three conventions for defining the zero position: 1. comfortable standing [34,35], 2. imposed position (e.g., subtalar neutral [36] or vertical tibia [37,38]), 3. referencing to the orientation of the underlying bony anatomy [39 43]. Nicholson et al. described variations in angles and segmental similarities for five different multisegmental foot models that the authors identified as being widely used in clinical and research settings [44]. The DuPont [45], Heidelberg [46], Oxford Child [47], Leardini (also known as Rizzoli) [48], and Utah [49] foot models were applied to data collected from ten unimpaired ambulators. Participants were instrumented with a custom marker set that included the necessary anatomic and technical markers for each of the five models. Kinematics of the hindfoot, forefoot, and hallux were assessed for variability between subjects, trials, and the instrumenting therapist (Fig. 4.5). The study identified consistent intra- and inter-therapist performance in marker placement and noted that differences in kinematics were primarily due to three factors: (1) markers used (and transformation from marker-based coordinate systems to segmentbased coordinate systems), (2) methods of angular calculations, and (3) definition of zero position. The definition of zero position also played a role in the significant offsets between the kinematic output of different models, and the authors noted that model- or lab-specific normative values would be required for clinical interpretation. Similar challenges in defining offsets have been reported previously, and the ability of these multisegmental foot

FIGURE 4.5 Average between therapist (PT), subject, and trial variability as shown by standard deviations across the gait cycle. On average, the Leardini model had the consistently highest variability in plantarflexion/dorsiflexion, and the Utah model had the highest variability with respect to forefoot ab/adduction. For most angles, models showed the highest variability during swing phase of gait. However, variability rarely changed by more than a few degrees throughout the gait cycle [44]. FF, Forefoot; Flex/Ext, flexion extension; HF, hindfoot; HX, hallux; Inv/Ev, inversion/eversion; IR/ER, internal rotation/external rotation; STD, standard deviation. Adapted from Nicholson et al [44].

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models to adequately capture deformities such as calcaneal valgus or collapsed longitudinal arch has been called into question [37,46]. As an alternative to establishing zero positions based on comfortable or imposed foot postures, there are numerous reports of multisegmental foot modeling methods that address motion at the level of the bony segments. Early work by Lundberg et al. reported on the kinematics of multiple foot segments using radio-opaque markers embedded in live subjects [39,41 43,50]. While these subjects underwent passive rotation under weightbearing conditions, roentgen stereophotogrammetry techniques were employed to measure motion in all three planes. A focus of this work was the contribution of each joint to out-of-plane motions. These works also provided verification that during motion, the talocrural joint axis moves continuously, with a rotation center near the midpoint of the bimalleolar axis. However, while these reports provided a robust and accurate measure of foot/ankle kinematics, the use of implanted markers and radiographs do not conform to the standards of noninvasive assessment and, as such, this technique will never be used clinically. These methods are reflected in development of the Milwaukee Foot Model (MFM), which calibrates neutral positions (zero positions) based on absolute bony alignment [39,40]. The MFM relies on measures of angles between adjacent bone segments as measured from weightbearing radiographs of the foot and ankle. Based on these measures from views in the A/P, mediolateral, and coronal planes, the orientation of bone-based axes in the global space is established. The weightbearing position is then replicated during motion capture using a tracing of foot position to duplicate the weightbearing posture. Following motion capture, the orientation of technical (marker-based) segments can be aligned with bone-based measures, and the resulting transformation matrix used to convert from technical axis systems to bone-based axis systems during walking trials. The MFM has been employed in numerous characterizations of adults and children with foot/ankle pathology [51 53], as well as assessments of post-operative change [54,55]. While its use of radiographs does pose challenges to a noninvasive approach to motion analysis, its use of bony alignment for neutral referencing represents a standard that cannot be attained by models that are entirely reliant on surface markers and fixed positions [56]. Regardless of zero position definition, most multisegmental models of the foot agree on the general morphology of the most common segments. Sagittal plane motion of the hindfoot tends to follow the three rockers of ankle motion. Whether it begins in a neutral or dorsiflexed position (depending on zero position), the hindfoot first flexes during load response, then goes through a period of progressive extension during stance, followed by flexion at toe-off, and then extension during swing. In the coronal plane, the hindfoot goes through progressive eversion throughout stance, followed by rapid inversion at toe-off as the foot transitions to a rigid lever for push-off. As reported by numerous models, forefoot motion tends to demonstrate minimal motion during load response. Once the foot achieves a plantigrade position, the forefoot generally extends throughout the stride, recovering into flexion following toe-off. Distinct shifts in coronal and transverse plan motion of the foot are also apparent at toe-off. For further discussion, please see the chapter on multisegment foot models.

4.6

Future areas of research

4.6.1 Biplane fluoroscopy The examination of foot and ankle motion using standard multisegmental kinematic models is associated with various rigid body assumptions and numerous technical limitations. As previously noted, the combined challenges of small bones, soft tissue artifact, and a lack of reliable anatomical landmarks lead to foot segments being grouped together when calculating kinematics. Additionally, the tracking of individual segments using skin-mounted markers has been reported to result in measurement errors ranging from 2.7 to 14.9 mm [57]. To work around the limitations imposed by tracking markers fixed to the skin surface, several groups have begun to employ biplane fluoroscopy systems to examine in vivo motion of anatomical structures of the foot and ankle during locomotion [58 63]. These approaches typically involve two X-ray sources and two image intensifiers spanning an activity volume (e.g., a walkway or treadmill). The subject’s foot and ankle can be volumetrically modeled via computer tomography (CT) scans that define the geometric edges of the bony structures of the foot, creating a calibration transform that can relate tracked images to the underlying bony geometry. Custom software can then be used to track bony rotations from the dynamic X-ray images, and the transform applied to generate true 3D kinematics at the individual joints of the foot and ankle. Current discussions on biplane fluoroscopic evaluation involve strategies to minimize radiation exposure and the establishment of consensus regarding strategies for establishing and tracking local coordinate axes on individual foot segments. Although fluoroscopic examination results in a relatively low dose of radiation to a distal extremity, the need for a CT scan (to customize the volumetric model to the individual’s unique anatomy) exposes the participant is to

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additional radiation. Strategies to address this potentially include low dose, weight bearing, cone beam CT scans [64] and magnetic resonance imaging [65,66] as replacements for the standard non-weight bearing CT scan. With regard to tracking standards, consensus has not been reached among investigators regarding strategies for attaching local coordinate axes to the segments being tracked and mechanisms for referencing motion of segments when calculating kinematics. Similar to marker-based modeling, investigators use different techniques during neutral referencing such as aligning talar and calcaneal coordinate axes parallel to the tibia [63] or establishing independent local coordinate axes based on individual anatomical features [67]. Efforts to identify benefits and drawbacks of each of these methods are ongoing. This technology is discussed in more detail in other chapters.

4.6.2 Modeling Significant progress has been made in understanding the kinematic profile of the multisegmental foot. Measurements of the joint kinetics associated with these movement patterns, however, remain limited. While kinematic analysis requires only the tracking of axis systems relevant to the underlying anatomy, kinetic analysis requires the definition of joint centers, the inertial properties of each segment of interest, and measures of segment stiffness. Additionally, the external force applied to each segment of interest must be isolated, and the ability to make these measurements is limited by the available technology. A force plate can provide highly accurate 3D measurements of an ambulator’s GRF, but the relatively large size of most commercially available plates limits their ability to isolate GRF measures to specific aspects of the foot. Early work in this area by Scott and Winter focused on separating the talocrural and subtalar joints into two individual articulations [68,69]. Their eight-segment model of the foot was built with monocentric single DOF joints, and plantar soft tissue was modeled as a nonlinear spring and nonlinear damper in parallel at seven independent sites. Unique to this model was its reliance on a decomposition of walking trials, in which only a portion of the subject’s foot struck the force plate during each walking trial. These methods were adapted in subsequent approaches to measuring multisegmental foot mechanics [70,71]. However, this novel approach to isolating GRF vectors to the individual segments also introduced high variability into the results, as well as the possibility for errors in estimates of joint rotation. Subsequent work in measuring multisegmental foot kinetics has sought new means of subdividing the single GRF measured by a force plate, distributing it across the relevant aspects of the foot. Pedobarography (plantar pressure assessment) presents one means of doing this, as pressure data that is synchronized with a step upon a force plate can be used to decompose the GRF vector and distribute it across the plantar surface. Numerous multisegmental models have used variations on this approach, with an assumption of proportionality used to distribute mediolateral and anteroposterior shearing forces in the same manner as the vertical pressure profile [34,72]. While the accuracy of the proportionality assumption has been called into question [73], its use has continued in recent reports that rely on center of pressure tracking to isolate the GRF to individual segments [74,75]. Further study of these approaches is ongoing, and with progress in the area of advanced imaging, opportunities to merge GRF data with tracking of true bony segments is no doubt forthcoming. With the advent of accessible biomechanical modeling software environments (e.g., SIMM, OpenSim, AnyBody), further work in foot modeling is also underway at the musculoskeletal level. Moving past the measurement of kinematics and kinetics associated with individual segments of the feet, these models also incorporate musculotendinous units that rely on imaging studies (e.g., CT scans) to identify anatomical origins and insertions for each muscle of interest [76 78]. While the technical characterization of these modeling approaches is beyond the scope of this chapter, their utility is readily apparent. Understanding the role of a patient’s underlying musculature in either causing or accommodating kinematic and kinetic abnormalities in gait will play a critical role in effectively planning treatment for that patient, regardless of whether that treatment is conservative (such as the proposal by Oosterwaal et al. to incorporate this approach into custom orthotics design [79]) or surgical in nature.

4.7

Conclusion

The closed chain system of loadbearing in the lower body during walking begins with the foot. While its structure makes up only 1.5% body weight, its role in ambulation is fundamental and critical. Its importance in this process highlights our need to understand its function, to accurately assess its mechanics, and (in the clinical care setting) to identify opportunities to remedy abnormalities. Improvements in motion capture technology over the past 30 years have served to better inform our ability to make these measurements; advances in imaging technology and modeling approaches will no doubt serve the same purpose in the years to come.

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J Pediatr Orthop B 2007;16(1):73 5. [20] Ristanis S, Stergiou N, Patras K, Tsepis E, Moraiti C, Georgoulis AD. Follow-up evaluation 2 years after ACL reconstruction with bonepatellar tendon-bone graft shows that excessive tibial rotation persists. Clin J Sport Med 2006;16(2):111 16. [21] Deltombe T, Detrembleur C, Hanson P, Gustin T. Selective tibial neurotomy in the treatment of spastic equinovarus foot: a 2-year follow-up of three cases. Am J Phys Med Rehabil 2006;85(1):82 8. [22] Molenaers G, Desloovere K, Fabry G, De Cock P. The effects of quantitative gait assessment and botulinum toxin a on musculoskeletal surgery in children with cerebral palsy. J Bone Jt Surg Am 2006;88(1):161 70. [23] Desloovere K, et al. How can push-off be preserved during use of an ankle foot orthosis in children with hemiplegia? A prospective controlled study. Gait Posture 2006;24(2):142 51. [24] Radtka SA, Oliveira GB, Lindstrom KE, Borders MD. The kinematic and kinetic effects of solid, hinged, and no ankle-foot orthoses on stair locomotion in healthy adults. Gait Posture 2006;24(2):211 18. [25] Chow DHK, Holmes AD, Lee CKL, Sin SW. The effect of prosthesis alignment on the symmetry of gait in subjects with unilateral transtibial amputation. Prosthet Orthot Int 2006;30(2):114 28. [26] Goujon H, et al. A functional evaluation of prosthetic foot kinematics during lower-limb amputee gait. Prosthet Orthot Int 2006;30(2):213 23. [27] Rankine L, Long J, Canseco K, Harris GF. Multisegmental foot modeling: a review," (in English). Crit Rev Biomed Eng 2008;36 (2 3):127 81. [28] Kadaba MP, Ramakrishnan HK, Wootten ME. Measurement of lower extremity kinematics during level walking. J Orthop Res 1990; 8(3):383 92. Available from: https://doi.org/10.1002/jor.1100080310. [29] Ballaz L, Plamondon S, Lemay M. Ankle range of motion is key to gait efficiency in adolescents with cerebral palsy. Clin Biomech 2010; 25(9):944 8. Available from: https://doi.org/10.1016/j.clinbiomech.2010.06.011. [30] Brodsky JW. Preliminary gait analysis results after posterior tibial tendon reconstruction: a prospective study. Foot Ankle Int 2004; 25(2):96 100. [31] Thomas R, Daniels TR, Parker K. Gait analysis and functional outcomes following ankle arthrodesis for isolated ankle arthritis. J Bone Jt Surg Am 2006;88(3):526 35. [32] Davis RB, Jameson E, Davids JR, Christopher LM, Rogozinski B, Anderson JP. The design, development, and initial evaluation of a multisegment foot model for routine clinical gait analysis. In: Harris GF, Smith PA, Marks RM, editors. Foot and Ankle Motion Analysis: Clinical Treatment and Technology. CRC Press; 2008. p. 425.

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[33] Nester CJ. Lessons from dynamic cadaver and invasive bone pin studies: do we know how the foot really moves during gait? J Foot Ankle Res 2009;2(1):18. Available from: https://doi.org/10.1186/1757-1146-2-18 2009/05/27. [34] MacWilliams BA, Cowley M, Nicholson DE. Foot kinematics and kinetics during adolescent gait. Gait Posture 2003;17(3):214 24. Available from: https://doi.org/10.1016/S0966-6362(02)00103-0 2003/06/01/. [35] Ringleb SI, Kavros SJ, Kotajarvi BR, Hansen DK, Kitaoka HB, Kaufman KR. Changes in gait associated with acute stage II posterior tibial tendon dysfunction. Gait Posture 2007;25(4):555 64. [36] Tome J, Nawoczenski DA, Flemister A, Houck J. Comparison of foot kinematics between subjects with posterior tibialis tendon dysfunction and healthy controls. J Orthop Sports Phys Ther 2006;36(9):635 44 Sep. [37] Carson MC, Harrington ME, Thompson N, O’Connor JJ, Theologis TN. Kinematic analysis of a multi-segment foot model for research and clinical applications: a repeatability analysis. J Biomech 2001;34(10):1299 307. [38] Woodburn J, Helliwell PS, Barker S. Three-dimensional kinematics at the ankle joint complex in rheumatoid arthritis patients with painful valgus deformity of the rearfoot. Rheumatol (Oxf) 2002;41(12):1406 12. [39] Kidder SM, Abuzzahab Jr. FS, Harris GF, Johnson JE. A system for the analysis of foot and ankle kinematics during gait. IEEE Trans Rehabil Eng 1996;4(1):25 32. [40] Long JT, Wang M, Harris GF. A model for the evaluation of lower extremity kinematics with integrated multisegmental foot motion. J Exp Clin Med 2011;3(5):239 44. [41] Lundberg A, Goldie I, Kalin B, Selvik G. Kinematics of the ankle/foot complex: plantarflexion and dorsiflexion. Foot Ankle 1989; 9(4):194 200 Feb. [42] Lundberg A, Svensson OK, Bylund C, Goldie I, Selvik G. Kinematics of the ankle/foot complex part 2: pronation and supination. Foot Ankle 1989;9(5):248 53 Apr. [43] Lundberg A, Svensson OK, Bylund C, Selvik G. Kinematics of the ankle/foot complex part 3: influence of leg rotation. Foot Ankle 1989; 9(6):304 9 Jun. [44] Nicholson K, et al. Comparison of three-dimensional multi-segmental foot models used in clinical gait laboratories. Gait Posture 2018;63:236 41. Available from: https://doi.org/10.1016/j.gaitpost.2018.05.013 Jun. [45] Henley J, Richards J, Hudson D, Church C, Coleman S, Kerstetter L. Reliability of a clinically practical multisegment foot marker set/model. Foot ankle motion analysis-clinical treat technol 2008;445 63. [46] Simon J, Doederlein L, McIntosh AS, Metaxiotis D, Bock HG, Wolf SI. The Heidelberg foot measurement method: development, description and assessment. Gait Posture 2006;23(4):411 24. Available from: https://doi.org/10.1016/j.gaitpost.2005.07.003 Jun. [47] Stebbins J, Harrington M, Thompson N, Zavatsky A, Theologis T. Repeatability of a model for measuring multi-segment foot kinematics in children. Gait Posture 2006;23(4):401 10. Available from: https://doi.org/10.1016/j.gaitpost.2005.03.002 Jun. [48] Leardini A, Benedetti MG, Berti L, Bettinelli D, Nativo R, Giannini S. Rear-foot, mid-foot and fore-foot motion during the stance phase of gait. Gait Posture 2007;25(3):453 62. Available from: https://doi.org/10.1016/j.gaitpost.2006.05.017 Mar. [49] Saraswat P, MacWilliams BA, Davis RB. A multi-segment foot model based on anatomically registered technical coordinate systems: method repeatability in pediatric feet. Gait Posture 2012;35(4):547 55. Available from: https://doi.org/10.1016/j.gaitpost.2011.11.022. [50] Lundberg A, Svensson OK, Nemeth G, Selvik G. The axis of rotation of the ankle joint. J Bone Jt Surg Br 1989;71(1):94 9 Jan. [51] Canseco K, et al. Distribution of segmental foot kinematics in patients with degenerative joint disease of the ankle. J Orthop Res 2018; 36(6):1739 46. Available from: https://doi.org/10.1002/jor.23807 Jun. [52] Kruger KM, et al. Segmental kinematic analysis of planovalgus feet during walking in children with cerebral palsy. Gait Posture 2017;54:277 83. Available from: https://doi.org/10.1016/j.gaitpost.2017.03.020. [53] Krzak JJ, et al. Kinematic foot types in youth with equinovarus secondary to hemiplegia. Gait Posture 2015;41(2):402 8. Available from: https://doi.org/10.1016/j.gaitpost.2014.10.027. [54] Canseco K, Long J, Smedberg T, Tarima S, Marks RM, Harris GF. Multisegmental foot and ankle motion analysis after hallux valgus surgery. Foot Ankle Int 2012;33(2):141 7. Available from: https://doi.org/10.3113/FAI.2012.0141. [55] Marks RM, Long JT, Ness ME, Khazzam M, Harris GF. Surgical reconstruction of posterior tibial tendon dysfunction: prospective comparison of flexor digitorum longus substitution combined with lateral column lengthening or medial displacement calcaneal osteotomy. Gait Posture 2009;29(1):17 22. Available from: https://doi.org/10.1016/j.gaitpost.2008.05.012. [56] Miller F. Cerebral palsy. New York: Springer-Verlag; 2005. [57] Maslen BA, Ackland TR. Radiographic study of skin displacement errors in the foot and ankle during standing. Clin Biomech 1994; 9(5):291 6. Available from: https://doi.org/10.1016/0268-0033(94)90041-8 Sep. [58] Campbell KJ, Wilson KJ, LaPrade RF, Clanton TO. Normative rearfoot motion during barefoot and shod walking using biplane fluoroscopy. Knee Surg Sports Traumatol Arthrosc 2016;24(4):1402 8. Available from: https://doi.org/10.1007/s00167-014-3084-4. [59] Cross JA, McHenry BD, Molthen R, Exten E, Schmidt TG, Harris GF. Biplane fluoroscopy for hindfoot motion analysis during gait: a modelbased evaluation. Med Eng Phys 2017;43:118 23. Available from: https://doi.org/10.1016/j.medengphy.2017.02.009. [60] Iaquinto JM, et al. Model-based tracking of the bones of the foot: a biplane fluoroscopy validation study. Comput Biol Med 2018;92:118 27. 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[62] Ito K, et al. Three-dimensional innate mobility of the human foot bones under axial loading using biplane x-ray fluoroscopy. R Soc Open Sci 2017;4(10):171086. Available from: https://doi.org/10.1098/rsos.171086 Oct. [63] Wang B, et al. Accuracy and feasibility of high-speed dual fluoroscopy and model-based tracking to measure in vivo ankle arthrokinematics. Gait Posture 2015;41(4):888 93. Available from: https://doi.org/10.1016/j.gaitpost.2015.03.008. [64] Linz C, et al. Performance of cone beam computed tomography in comparison to conventional imaging techniques for the detection of bone invasion in oral cancer. Int J Oral Maxillofac Surg 2015;44(1):8 15. Available from: https://doi.org/10.1016/j.ijom.2014.07.023. [65] Akbari-Shandiz M, Lawrence RL, Ellingson AM, Johnson CP, Zhao KD, Ludewig PM. MRI vs CT-based 2D 3D auto-registration accuracy for quantifying shoulder motion using biplane video-radiography. J Biomech 2018;. Available from: https://doi.org/10.1016/j.jbiomech. 2018.09.019 2018/09/29/. [66] Moro-oka TA, et al. Can magnetic resonance imaging-derived bone models be used for accurate motion measurement with single-plane threedimensional shape registration? J Orthop Res 2007;25(7):867 72. Available from: https://doi.org/10.1002/jor.20355 Jul. [67] Kruger K. et al. Development of an anatomic coordinate system to calculate hindfoot kinematics using biplane fluoroscopy. In: 2018 Annual meeting of the Gait and Clinical Movement Analysis Society, Indianapolis, IN, May 22 25, 2018; 2018. [68] Scott SH, Winter DA. Talocrural and talocalcaneal joint kinematics and kinetics during the stance phase of walking. J Biomech 1991; 24(8):743 52. [69] Scott SH, Winter DA. Biomechanical model of the human foot: kinematics and kinetics during the stance phase of walking. J Biomech 1993; 26(9):1091 104. [70] Bruening DA, Cooney KM, Buczek FL. Analysis of a kinetic multi-segment foot model. Part I: model repeatability and kinematic validity. Gait Posture 2012;35(4):529 34. Available from: https://doi.org/10.1016/j.gaitpost.2011.10.363. [71] Bruening DA, Cooney KM, Buczek FL. Analysis of a kinetic multi-segment foot model part II: kinetics and clinical implications. Gait Posture 2012;35(4):535 40. Available from: https://doi.org/10.1016/j.gaitpost.2011.11.012. [72] Saraswat P, MacWilliams BA, Davis RB, D’Astous JL. Kinematics and kinetics of normal and planovalgus feet during walking (in English). Gait Posture 2014;39(1):339 45. Available from: https://doi.org/10.1016/j.gaitpost.2013.08.003 Article. [73] Yavuz M, Botek G, Davis BL. Plantar shear stress distributions: comparing actual and predicted frictional forces at the foot ground interface. J Biomech 2007;40(13):3045 9. Available from: https://doi.org/10.1016/j.jbiomech.2007.02.006 2007/01/01/. [74] Bruening DA, Cooney KM, Buczek FL, Richards JG. Measured and estimated ground reaction forces for multi-segment foot models. J Biomech 2010;43(16):3222 6. Available from: https://doi.org/10.1016/j.jbiomech.2010.08.003 2010/12/01/. [75] Bruening DA, Takahashi KZ. Partitioning ground reaction forces for multi-segment foot joint kinetics (in English). Gait Posture 2018;62:111 16. Available from: https://doi.org/10.1016/j.gaitpost.2018.03.001 Article. [76] Malaquias TM, et al. Extended foot-ankle musculoskeletal models for application in movement analysis. Comput Methods Biomech Biomed Engin 2017;20(2):153 9. Available from: https://doi.org/10.1080/10255842.2016.1206533. [77] Saraswat P, Andersen MS, Macwilliams BA. A musculoskeletal foot model for clinical gait analysis. J Biomech 2010;43(9):1645 52. Available from: https://doi.org/10.1016/j.jbiomech.2010.03.005 Jun 18. [78] Qian Z, Ren L, Ren L, Boonpratatong A. A three-dimensional finite element musculoskeletal model of the human foot complex. In: Lim CT, Goh, JCH, editors. 6th World congress of biomechanics (WCB 2010), Singapore, August 1 6, 2010, vol. 31. Berlin, Heidelberg: Springer. In: IFMBE Proceedings; 2010. [79] Oosterwaal M, et al. Generation of subject-specific, dynamic, multisegment ankle and foot models to improve orthotic design: a feasibility study. BMC Musculoskelet Disord 2011;12:256. Available from: https://doi.org/10.1186/1471-2474-12-256.

Chapter 5

Bone, Cartilage, and Joint Function Michael T. Perez1 and Jennifer S. Wayne1,2 1

Department of Biomedical Engineering, Virginia Commonwealth University, Richmond, VA, United States, 2Department of Biomedical Engineering

& Mechanics, Virginia Tech, Blacksburg, VA, United States

Abstract The human foot comprises 28 individual skeletal bones and more than 40 articulations. These bones and articulations bear the body’s weight and permit individual joint movements to accommodate a variety of activities and terrains. Achieving these functional demands requires resilient bone, articular cartilage, and other soft tissue structures that can withstand millions of loading cycles each year throughout an individual’s life. This chapter focuses on bone, cartilage, and joint function in general and then with specific references to the functions in the foot.

5.1

Bone components and structure

While the anatomy of skeletal bones varies considerably both in the human body and in the foot, the fundamental components of each remain the same—solid components hydrated by water. By volume, the solid extracellular matrix (ECM) makes up the largest percentage. This matrix is dominated by the inorganic mineral hydroxyapatite (HA) (65% 70% dry weight) and the organic protein Type I collagen (25% 30%) [1]. HA primarily provides the bone’s strength, while the network of collagen fibers reinforces against the brittleness common in crystalline only materials and increases bone’s toughness, analogous to other composite structures. Lastly, water comprises approximately 25% of the total weight of bone [2]. At the macroscopic level, skeletal bone is classified as either cortical bone or cancellous, each of which have specific structural arrangements and functions. Other names for cortical bone are compact or dense bone, while other names for cancellous bone are porous, spongy, or trabecular bone. Cortical bone is organized into tightly packed osteons which are circular structural units containing a central canal for blood supply and innervation. By comparison, cancellous bone is organized into an open cell latticework of units called trabeculae, which are commonly described as plate or rod shaped. The bones in the foot follow the same internal structural organization as bones in the rest of the body, with dense cortical bone forming a protective layer around the interior spongy cancellous bone [3] (Fig. 5.1). Like other living tissues, cells are interspersed throughout bone for maintenance of the ECM as well as to remodel or repair as needed based on local use and loading. The metabolic activity of bone cells helps in repair and maintenance of bone tissue. The osteoclasts and osteoblasts control ECM resorption and formation, respectively. Osteocytes help maintain the ECM and play a major role in cell signaling by sensing the bone’s mechanical environment through mechanosensory mechanisms to regulate osteoclast and osteoblast activity. Under normal circumstances, these cell types work synergistically to repair acute injuries and continuously remodel as needed. In cases of repair, osteoblasts first form woven bone. Woven bone is immature bone that is quickly produced but mechanically weak. Over time, woven bone is remodeled and replaced by the mechanically stronger lamellar bone. The remodeling activity in bone is heavily regulated by stresses, strains, and micro fractures that the osteocytes experience. As a result, the repeated loading of bones in the foot during daily activities provides significant cues for bone remodeling [5]. The risk of stress fractures in the bones are higher when the repeated loading is in tension while the risk is usually lower for compressive loading, assuming there are no muscular or metabolic pathologies at play [6]. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00028-7 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 5.1 Diagram representing cortical and trabecular bone organization, here in the femur [4]. Reprinted with permission Hayes WC. Biomechanics of cortical and trabecular bone—implications for assessment of fracture risk. In: Mow VC, Hayes WC, editors. Basic orthopaedic biomechanics. 1st ed. Wolters Kluwer; 1991.

Bone biomechanical behavior requires understanding two key concepts. The first is the mechanical behavior of bone’s materials. Cortical bone is organized into parallel oriented osteons consisting of lamellar bone [7]. Similarly, the plate and rod shaped trabecular bone also show directional orientation. Remodeling of cortical and cancellous bone can induce changes to this preferred orientation in the bone’s structure dependent on mechanical cues. Bone’s oriented microstructure introduces the important characteristic of anisotropy. Thus, the behavior of the bone is strongly dependent on the direction of loading relative to the alignment of its structure, indicating bone behaves anisotropically, characteristically as an orthotropic material. Bone shows the greatest stiffness when loaded longitudinal to the bone’s osteon alignment, while loads transverse to this alignment show lower stiffness. This anisotropic behavior makes bone a more complex structure and gives insights into whole bone biomechanics [8]. The microstructure of a bone can vary quite significantly between different bones and even between the locations within a specific bone (Fig. 5.2). For example, the posterior calcaneus has trabeculae oriented from the talus to the heel’s ground contact point, whereas the trabeculae are oriented in a very different manner in the calcaneal neck. Meanwhile, the metatarsals have less organized structure in their base and head but a highly aligned section in their diaphysis. The varied microstructures impact the local mechanical behavior of the different regions of the bone. At this material level, bone is often modeled as linearly elastic with some viscoelastic behavior when hydrated. This viscoelasticity defines a strain rate dependence which affects both the elastic modulus and ultimate strength, whereby higher strain rates increase both properties [8]. The second key concept is how the different structural elements combine to influence skeletal bone behavior as a whole. When reviewing whole bone biomechanics, the cortical and cancellous bones each have their own role to play.

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FIGURE 5.2 Sagittal view from a radiograph of a right foot where trabecular orientation is visualized, particularly in the calcaneus and talus. Image courtesy VA Puget Sound, Seattle WA.

The compact, highly aligned osteons of cortical bone comprises the outer shell of the tarsal bones and are important in providing strength to the shaft (diaphysis) of the metatarsals. The cancellous interior of bones, while having lower strength and rigidity than cortical bone, is important to load transmission at the joints and its ability to dampen sudden stresses. Together, these different types of bones provide the strength and flexibility needed to withstand the stresses of daily activities [8]. To quantify local characteristics of a skeletal bone, different parameters/measurements are reported. Bone volume fraction and cortical and trabecular thickness are measures that can be used to quantify a bone’s morphology. Meanwhile, the bone mineral density (BMD) and degree of anisotropy can inform on the loading pattern of the bone. These measurements explain 98% of the variation in bone’s elastic properties [9]. Bone volume fraction (BV/TV) is the volume of mineralized bone divided by the total volume in the bone sample. This is an important parameter in trabecular bone as it has the latticework of mineralized bone with voids throughout. Many computed tomography (CT) and micro-CT scanners can calculate these values, but variations in technique exist. For CT scanners incapable of resolving individual trabeculae (scan resolution larger than B80 µm) an estimate can be made whereby BV/TV is derived from the measured trabecular bone density divided by 1200 mg HA/cm3 (assumed density of fully mineralized bone) [10,11]. In the foot, bone volume fractions correlate with lifestyle and personal traits that introduce stresses which induce bone remodeling to best withstand the stresses, specifically in the calcaneus. For example, the body weights of groups of nonrunners explain B80% of the variation seen in BV/TV, trabecular thickness, and trabecular density for this population. Meanwhile, endurance runners do not show this trend. Instead, the age at onset of running better explained the variation in BV/TV [10]. These two trends help highlight the complexity of factors that can change the morphology of bones and by extension, their physical properties. In one case the variation of the daily load (body weight) changed the density of the bone, while the other case, the earlier the individual begins high cyclic loading impacts the density of the bone later in life. Thus bone remodeling response may change with aging while also change with the current loads experienced. Another quantitative measure from dual energy X-ray absorptiometry (DEXA) or CT imaging is BMD. DEXA uses two low dose X-rays and measures the attenuation of the signals passing through the body. This is the common clinical procedure for identifying low BMD which is a key sign of osteoporosis [12]. As with DEXA, BMD can be calculated from a CT scan for a given region in the bone based on the attenuation of the CT signal. Various CT algorithms are used to determine BMD for both cortical bone and cancellous bone where marrow tissue is mixed into the voids in the bone [13]. Either procedure can arrive at similar BMD; however, different biases in the procedure cause CT scanderived measurements to be lower than the gold standard DEXA measurements [14]. BMD measurements have been used to help inform locations in bone that would best resist pullout of cortical screws used in internal fixation for arthrodesis or for fracture repair. Research in this area has provided detailed mapping of where the cortical shell and cancellous bone is the thickest and has the highest mineral density. For example, in the medial cuneiform researchers have shown the portions of the bone with the highest BMD to guide where cortical screws may have the best purchase and thus survival rates [15,16]. Other investigators have suggested using BMD in creating injury probability curves for bones in the foot and ankle [17].

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Another common way to understand bone function is by characterizing trabecular anisotropy. Several techniques exist to evaluate this for a bone region (mean-intercept length, star volume distribution, and star length distribution methods), but they all capture the variation of the trabeculae in the different directions [18]. The trabeculae in the bones of the foot, like those in the femoral head, orient themselves along the direction of loading, which can be visibly seen in imaging scans as indicated previously (Fig. 5.2). In the foot, areas of high anisotropy include the posterior calcaneus (posterior talar articular surface to the calcaneal tuberosity) and the talar neck. These regions transmit body weight forces to the ground or midfoot, respectively [19]. Morphological studies of fossil records for trabecular alignment have helped determine the likelihood of bipedal movement by members of the Hominidae family as this induces different loading patterns on the foot [20,21]. Other quantitative approaches to understand bone architecture are based on measuring trabecular thickness, number, and separation. These measurements in the talus have been used to inform on the different means of locomotion in fossil records [22]. These sets of trabecular bone measurements better represent a specimen’s types and levels of activities [23]. In the calcaneus, variation in the ultimate strength of the cancellous bone is highly correlated with these trabecular measurements. Some studies have shown trabecular measurements to be a better predictor of ultimate strain than BMD [24]. A drawback is that accurate values for these measurements require a research grade micro-CT. While clinical high resolution peripheral quantitative CT scans can be used to derive values for thickness, number, and separation of trabeculae, these procedures result in systematic overestimation of trabecular numbers and underestimations of trabecular thickness and spacing [25]. Decreases in the mechanical cues due to stress shielding or decreased activity will lead to increased bone resorption by osteoclasts. Other factors that can increase bone resorption or decrease bone production are diet, hormones, and medications like glucocorticoids. The loss of bone leads to osteoporosis and bones become more fragile and will fracture at lower loads. In the foot and ankle, osteoporosis fractures are rarer than the hip or wrist, but transient osteoporosis may occur. Often times there are signs of bone loss and swelling at the site of pain. However, rest and conservative treatment are often all that is needed. There are rare cases where the affected area “migrates” within the foot but it also appears these cases are recoverable with conservative treatment [26,27].

5.2

Cartilage

Articular cartilage is the soft connective tissue that covers the regions of a bone’s surface that articulate with other bones, i.e., in joints. Articular cartilage serves two main functional purposes in joints. It helps to lubricate articulating surfaces thereby ensuring low levels of friction to enable smooth joint movement and reduce wear; and it helps distribute loads over the joint surface through cartilage deformation and fluid pressurization, thereby reducing contact stresses [28]. Articular cartilage is composed of a mixture of components that enable its functions. It is highly hydrated, with 70% 85% water by total weight. Uninjured cartilage has very low cellularization (chondrocytes, 10% of volume) so the majority of the tissue is an ECM composed of proteoglycan aggregates and Type II collagen. Proteoglycans are large, negatively charged macromolecules with chains of different glycosaminoglycans (GAGs). The electrostatic repulsive forces of the GAGs promote a hydrophilic environment and resist compression of the tissue. The collagen component serves to resist tension experienced by the tissue. Cartilage ECM is often described as an intertwining porous mesh of collagen fibrils and proteoglycan aggregates through which fluid flows (Fig. 5.3). The combination of a high water content, the intrinsic nature of the ECM, and the pressurization and flow of fluid induces a viscoelastic response of the tissue to mechanical loading. Through this response, articular cartilage can withstand high cyclic compression loading, acting like a shock absorber. When cartilage is injured or diseased, its ability to deform, pressurize the interstitial fluid, and transfer load is altered which may lead to higher stress in the subchondral bone [29]. Articular cartilage lacks a blood supply, lymphatic system, and nerves, which led many to believe that the tissue lacked any biological activity. In the second half of the 20th century, cartilage was revealed to be a more complex tissue than previously believed. The presence of chondrocytes, cells within the tissue matrix that produce and maintain the cartilage ECM, was demonstrated, and the tissue was revealed to have important structural organization through its thickness despite being relatively thin (B1 to 5 mm, depending on the joint). Three main zones were documented (Fig. 5.4). The topmost is the superficial tangential zone (STZ) with highly organized collagen fibrils tangential to the articulating surface, very few proteoglycans, and flattened/elongated chondrocytes. The middle zone comprises randomly oriented collagen interspersed with proteoglycans and spherical chondrocytes. The deep zone has radially oriented collagen that crosses the tidemark and fewer proteoglycans with columnar oriented chondrocytes. With the discovery of chondrocytes in cartilage, the idea of regenerating the tissue has intrigued the scientific and clinical communities. To help with this work, quantifying the biomechanical function of articular cartilage became of

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FIGURE 5.3 Schematic of the intertwining mesh concept of the solid collagen fibrils/proteoglycan aggregates in articular cartilage [28]. Reprinted with permission Mow VC, Proctor CS, Kelly MA. Biomechanics of articular cartilage. In: Nordin M, Frankel VH, editors. Basic biomechanics of the musculoskeletal system. 2nd ed. Wolters Kluwer; 1989.

FIGURE 5.4 Structure of cartilage through its thickness showing cell distribution (A: histological section, B: schematic) and collagen architecture (C: schematic, D: scanning electron micrographs) [30]. Reprinted with permission: Buckwalter JA, Mow VC, Ratcliffe A. Restoration of injured or degenerated articular cartilage. J Am Acad Orthop Surg 1994;2:192 201. https://insights.ovid.com/pubmed?pmid 5 10709009.

paramount importance. From this work arose several different structural constitutive models with associated equations representing the behavior of different components in the tissue [31]. One of the first models to rise to prominence was the biphasic theory. The theory built on previous work by Mow, Torzilli, and Mansour and treated the tissue as having two components, a linearly elastic ECM and a viscous fluid [32,33]. The theory was first used to predict creep and stress relaxation of cartilage as governed by the diffusional drag from the interstitial fluid moving through the solid matrix [34]. The role that each zone within the tissue plays in the function of articular cartilage is evident from a biphasic point of view. The STZ with few proteoglycans is compacted upon the compressive stresses from joint articulation, making

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the already dense collagen layer even more compact. The result is a restriction of fluid flow out of the tissue, effectively pressurizing the fluid within the tissue and reducing the stresses the solid matrix endures. The middle zone with its high proteoglycan content is critical to withstanding the compressive stresses from joint forces. The deep zone with its radially oriented fibers serves to anchor the tissue to the underlying bone. While the biphasic theory was a major advance in our understanding and description of articular cartilage response to loading, concerns arose over its ability to properly reflect the tensile characteristics of the ECM, specifically, the collagen network, as well as whether the ECM was viscoelastic instead of purely elastic. Updates and additions to this model have been implemented over the years to better characterize different structures/orientations in the tissue. To better reflect tissue behavior in tension, an updated model considered the ECM to be a network of collagen fibrils that can only resist tension, surrounded by a hydrophilic gel of proteoglycans to resist compression [35]. Still, the heterogeneity in matrix composition and fibril network orientation was difficult to adequately replicate. Another extension to the biphasic model was the triphasic model which incorporated a third component to account for the role of mobile ions interacting with the fixed negatively charged glycosaminoglycan chains. The electrostatic repulsive forces between the GAG chains are altered by the presence of ions. These mobile ions can shield the fixed charged protein chains and reduce their repulsive force, which impacts the mechanical properties of the cartilage [36]. This additional component permitted the model to explain both the mechanical and chemical loads experienced in the cartilage. Many other improvements have been made, e.g., to incorporate the intrinsic viscoelastic nature of the ECM, or a bilinear tension/compression constitutive equation for the ECM, all designed to expand our understanding and predictive ability for cartilage function [37,38]. In the foot, many bones articulate with three or more separate joints. Thus, several bones, like the cuneiforms and talus, have a large portion of their surface covered in articular cartilage. In addition to the large number of articulating surfaces, some of the surfaces themselves are relatively large due to the range of motion exhibited by the individual joints. For instance, the metatarsal heads have a large region of cartilage to allow full range of motion of the metatarsophalangeal joint in the sagittal plane [39]. The talar dome also has a large cartilage coverage to allow rotation about the ankle (talocrural) joint (Fig. 5.5). In healthy individuals, cartilage thickness ranges from 0.6 to 1.2 mm depending on location [40,41]. Cartilage on the talus and distal tibia show higher cartilage thickness while values decrease in the midfoot between the cuneiforms and navicular [41]. The cartilage of the talar dome has been compared with other major joints in the leg, and there are two notable differences. First, the concentration of proteoglycans and water in the cartilage of the ankle joint is higher. Second, there is a higher rate of proteoglycan break down and synthesis at the ankle. These differences produce a stiffer cartilage in the ankle that has a lower permeability. Besides differences in the mechanical properties of the cartilage, the chondrocytes in the ankle also show some noticeable differences in healing. After the formation of a lesion, chondrocytes in the knee downregulate collagen and proteoglycan production while ankle chondrocytes upregulate the production of both [42]. The ankle joint has a lower rate of osteoarthritis (OA) than the knee. A possible reason for this is the behavior of the subchondral bone. In joints that commonly develop OA, cartilage degradation is accompanied by an increased density of subchondral bone. In the ankle, when there are signs of cartilage degradation, the subchondral bone shows signs of bone density decrease. This makes it possible that the higher rate of ECM production by chondrocytes and decreased subchondral bone density helps prevent progressive OA in the ankle [42]. While the ankle has signs that it is better at resisting OA than the knee, the cases where OA does develop have often experienced previous trauma [42,43]. Other areas of the foot’s cartilage can also develop OA after injuries.

FIGURE 5.5 Artistic rendering of a right talus depicting the large surfaces for articulation with the distal tibia/fibula and navicular.

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Radiographs of the midfoot after dislocations show many patients have developed OA in joints around their injury site [44]. Other than direct injuries, mid and hindfoot OA can develop when the ipsilateral ankle undergoes arthrodesis [45]. This has been hypothesized to be a result of irregular movement and loading in the ipsilateral foot when walking to compensate for the fused ankle.

5.3

Joint functions

The most basic definition of a joint is the location where two bones contact one another and allow an amount of relative movement between the two. The study of joints also incorporates the study of the soft tissues that allow or prevent different motions in the joint. These include the ligaments that constrain a set range of motion, muscles that generate forces for motion, and cartilage, the role of which has already been discussed. Several of the joints in the foot have added complexity, for example those that consist of multiple bones and points of contact making up what is considered a single joint. The unique geometry and role of the soft tissues contribute to the degrees of freedom and kinematic motion of each joint. By identifying the associated motions and forces, the contribution of each joint to the biomechanics of the foot as a whole is better understood. In turn, understanding the injuries and points of failure in these joints help understand the effects and shortcomings of different treatments for the foot.

5.3.1 Talocrural joint The talocrural or ankle joint is composed of the contact areas between the distal tibia/fibula and the talar dome and is thus also called the tibiotalar joint (Fig. 5.6). The talar body sits between the malleoli of the distal tibia and fibula giving rise to the woodworking analogy of a mortise and tenon joint. This anatomical congruity of the three bones confers substantial stability to the joint. A large portion of the talar dome and body is covered with cartilage because the tibial plafond articulates with the superior surface of the talus while the malleoli articulate on the medial and lateral sides of the talar body. The joint is surrounded by a synovial capsule and several ligament complexes to constrain the range of motion. The talocrural joint largely functions as a hinge as the talus rotates between the malleoli enabling the motions of dorsiflexion (pointing toes upward) and plantarflexion (pointing toes toward the floor). At different angles of plantarflexion and dorsiflexion, there are different amounts of lateral translation and inversion/eversion permitted in the joint. These are largely limited by the major collateral and deltoid ligaments [39]. The additional degrees of freedom in the joint cause the axis of rotation to shift superiorly on the medial side of the ankle during dorsiflexion [46]. While the axis of rotation is dependent on the flexion angle, the average flexion axis of rotation is angled medial-anterior to lateral-posterior when viewed in the transverse plane (Fig. 5.7). This angle is a result of the medial malleolus being more anteriorly located than the lateral malleolus so the talar dome is not perfectly oriented anterior to posterior. In motion analysis studies, the axis is most often calculated based on the kinematic motion during walking of key landmarks on the subject. However, some research creates the axis without kinematic data and just uses the morphology of the talus from CT imaging [47]. The drawback is that this results in a single axis of rotation over the whole of flexion motion. An additional feature of the talocrural joint is that the width of the articulating surface on the talus is larger anteriorly than posteriorly. During dorsiflexion, this causes the mortise to widen between the tibial and fibular malleoli

FIGURE 5.6 Anterior view of a right talocrural (i.e., tibiotalar or ankle) joint, comprised of the distal tibia and fibula surrounding the talus.

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FIGURE 5.7 Superior view of right foot bones with the malleoli (blue bone segments) shown on either side of the talus and the talocrural axis (dashed line).

which stretches the interosseous membrane and other soft tissues between the tibia and fibula. In plantarflexion, the narrower talar surface does not fit as tightly in the mortise formed by the malleoli so some translation is possible. This introduces instability in the joint during plantarflexion, allowing some degree of inversion and eversion and requiring the ligaments to provide significant support. The lateral collateral ligaments and deltoid ligaments minimize inversion and eversion, respectively. These two ligamentous structures also play a role in resisting excessive flexion angles. The lateral collateral ligament bands are observed to be strained throughout the range of motion. However, the deltoid ligament’s anterior bands become taut during plantarflexion, and the posterior bands become taut during dorsiflexion [39]. The most common injury to the talocrural joint in athletes is ligament sprain [48]. These injuries are often caused by excessive inversion or eversion of the joint while the foot is plantarflexed. This position places large amounts of stress on the ligaments since the bones of the joint, as previously mentioned, do not provide as much stability in plantarflexion [39]. Depending on severity of the injury, conservative treatment may be sufficient to treat the symptoms of this injury by reducing the swelling and pain from the injury; however, chronic instability in the joint can result. Other injuries may degrade the cartilage of the talar dome, but the unique responses of the tissue to damage discussed earlier may reduce the risk of these cases progressing to OA [40].

5.3.2 Talocalcaneal (subtalar) joint The subtalar joint is the articulation between the talus and the calcaneus. This joint is responsible for most of the inversion and eversion motion in the foot. A unique consideration for this joint is that these two bones actually have two distinct articulating surfaces each with their own synovial capsules (Fig. 5.8) [49]. The posterior articulating facet is an interaction between the convex superior portion of the calcaneus against the concave portion of the inferior posterior talus. This is sometimes considered the only aspect of the subtalar joint. The second articular facet is located between the inferior portion of the talar neck and head and the superior portion of the anterior calcaneus [39]. It is sometimes considered part of the “talocalcaneonavicular joint” since it shares a synovial cavity with the talonavicular joint. Since the subtalar joint is the source of most of the inversion and eversion motion in the foot, one would expect the axis of rotation to run anterior to posterior. Instead, the axis runs antero-medio-superior to postero-latero-inferior and the axis of rotation is described as a helical or screw axis. This axis can be approximated by circumscribing spheres to fit the two articular surfaces on the talus and connecting the centers of the spheres (Fig. 5.9). Similar to the talocrural joint, the axis changes orientation over the joint’s range of motion but this has not been well characterized [50]. Some of the crucial ligaments for the talocrural joint also play a role in stabilizing and controlling the subtalar joint. This results from portions of the collateral ligaments and deltoid ligament inserting on the tibia or fibula, crossing the talus without inserting, and terminating on the calcaneus. These provide more resistance to excessive motion than the two ligament complexes connecting the calcaneus and talus directly [39]: the interosseous and lateral talocalcaneal

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FIGURE 5.8 (Top) Medial view of the subtalar joint formed by articulations between the talus and calcaneus of a right foot. (Bottom) Superior view of the calcaneus highlighting the articular surfaces of the subtalar joint.

FIGURE 5.9 (A) Inferior view of the talus with colored regions depicting articulations; (B) Sphere fits to the articular surfaces on the inferior talus; (C) Inclined lateral view of the talus with sphere fits from (B) superimposed; (D) Axis of subtalar rotation connecting centers of spheres represented on the talus [49]. Reproduced under Creative Commons Attribution License: Parr WC, Chatterjee HJ, Soligo C. Calculating the axes of rotation for the subtalar and talocrural joints using 3D bone reconstruction. J Biomech 2012;45(6):1103 7. https://doi.org/10.1016/j.jbiomech.2012.01.011.

ligaments. Besides the ligaments, the subtalar joint is supported by the tendons that wrap around the calcaneus before inserting on other bones in the inferior (plantar) anterior portion of the foot. These tendons include the peroneus longus, peroneus brevis, flexor hallucis longus, flexor digitorum longus, and tibialis posterior (Fig. 5.10).

5.3.3 Transverse tarsal joint The transverse tarsal joint, also referred to as the Chopart’s joint, comprises the talonavicular joint and the calcaneocuboid joint (Fig. 5.11). These joints serve an important role in providing flexibility in foot but can also form a rigid lever during the heel off phase of gait. This joint is also a location for one of the levels of within-foot amputation (midtarsal). The talonavicular joint is formed by the anterior portion of the talar head contacting the posterior surface of the navicular and is constrained by an array of ligamentous structures. The cartilage on the talar head extends to the inferior

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FIGURE 5.10 Medial view of the right foot showing course of tendons [51]. Credit BodyParts3D, Copyright 2008 Life Science Integrated Database Center licensed by CCView-Inheritance 2.1 Japan.

FIGURE 5.11 Anterior view of the hindfoot of a right foot showing the articulating surfaces on the talus and calcaneus making up the transverse tarsal joint.

surface of the talus where the cartilage becomes part of the anterior aspect of the subtalar joint [52]. The head of the talus is convex and the navicular articulation only covers a small portion of the head, suggesting that this would be an unstable ball and socket joint. Stability comes from the bifurcate ligament on the lateral side and the inferomedial and superomedial calcaneonavicular ligament. The tibialis posterior and the flexor tendons also play a dynamic role in stabilizing the joint [53]. The calcaneocuboid joint is a saddle joint formed from the anterior surface of the calcaneus contacting the posterior surface of the cuboid. The range of motion in the calcaneocuboid joint is more restricted than the talonavicular joint as a result of adjacent bony contact. The posterior surface of the cuboid has an inferomedial bone projection. This projection sits in a fossa of the calcaneus during inversion of the foot. The superomedial portion of the calcaneus also has a bony ridge that overhangs the cuboid to prevent dislocation of the cuboid superiorly [52]. The talonavicular and calcaneocuboid joints are directly impacted by the inversion and eversion of the subtalar joint. Changes in the subtalar joint determine whether the transverse tarsal joint forms a rigid or flexible connection from the hindfoot to the rest of the foot. The transverse tarsal joint changing between a flexible and a rigid joint is known as the midfoot locking mechanism and plays an important role in the gait cycle. During normal gait, as the foot contacts the ground, there is an initial eversion of the subtalar joint. This motion brings the axes of the talonavicular and

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calcaneocuboid joints into the same alignment. When this occurs, the midfoot is more flexible and free to move and accept variations in the contacting terrain. From the stance phase to heel off phase, the subtalar joint experiences inversion, which moves the axes of the talonavicular and calcaneocuboid joints out of alignment. This causes locking of the midfoot into a rigid lever that is used to push the leg forward through the toe off phase [54]. Disease or injuries throughout the foot can impact the midfoot locking mechanism. Calcaneal fractures or hindfoot arthrodesis impact the behavior of the subtalar joint. This changes the kinematics of the transverse tarsal joint and may lead to increases in stress or ligamentous injuries if the midfoot locking mechanism fails during the gait cycle [52]. This illustrates how changes to a portion of the foot can have meaningful impact on the function and biomechanics of the foot as a whole. Another disease state that impacts the transverse tarsal joint is flatfoot deformity. The tibialis posterior has insertion sites on the navicular and plays an important role in inversion of the foot. When there is an insufficiency of the muscle or tendon, subtalar inversion cannot properly occur [55]. This means the axes of the talonavicular and calcaneocuboid will remain in alignment during the gait cycle and prevent proper locking of the midfoot.

5.3.4 Tarsometatarsal joint The tarsometatarsal joint, also called Lisfranc’s joint, is a series of joints between the distal surfaces of the three cuneiforms and cuboid with the proximal bases of the five metatarsals (Fig. 5.12). The medial three metatarsals each articulate on the distal surface of one of the cuneiforms while the lateral two metatarsals articulate on the cuboid. The articular surfaces of the tarsometatarsal joint are mostly flat, relying on dense ligament structures between the metatarsal and tarsal bones to prevent most motion in the joint. A unique portion of the joint’s anatomy is around the intermediary cuneiform and the second metatarsal. The distal surface of the intermediary cuneiform is recessed compared to the cuneiforms on either side of it. This means the second metatarsal base sits in a mortise between the medial and lateral cuneiforms. The recessed second metatarsal base has several bands of a ligament that connects it to the medial cuneiform. This ligament is important because there is no proximal ligament connection between the first and second metatarsals, therefore this ligament plays a major role in preventing separation of the first and second rays of the foot [56]. The tarsometatarsal joint is the connection point between the metatarsals contacting the ground and the rest of the foot. Because there is limited motion in the joint, the kinematics and forces of the joint have not been a major focus of research. This likely explains why treatment of injuries to the proximal bases of the metatarsals and ligament injuries have poor surgical outcomes and high rates of arthritis after injury [57,58]. With a better understanding of this joint’s mechanics, there is hope injury treatments can be improved.

5.3.5 Metatarsophalangeal joint The metatarsophalangeal joint consists of the articular surfaces between each of the metatarsal heads and the proximal phalanges of each of the five toes. The metatarsal heads are covered in cartilage from the plantar to dorsal surfaces. This large cartilage articulation supports the low friction movement of the phalanges and metatarsals throughout their considerable range of motion (up to 75 degrees extension and 40 degrees flexion). All five metatarsal heads have a convex surface that contacts the concavity of the proximal phalanges. These joints largely function as hinges as the phalanges rotate about the head of the metatarsals. When viewed in anterior-posterior radiographs, it is important to note that the long axes of the metatarsal and proximal phalanges in

FIGURE 5.12 Closeup view of the dorsal midfoot showing the tarsometatarsal joint (red line) and the second metatarsal (MT2) sitting in the mortise formed by the cuneiforms (CN1 3).

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each joint do not align. In the first ray, the angle between the long axis of the metatarsal and proximal phalanges is normally less than 15 degrees. One of the most common injuries/disease states in this joint occur when this halluxmetatarsophalangeal angle exceeds 15 degrees. This is called hallux valgus but may be more commonly known as a bunion since the metatarsal head irritates the medial side of the foot.

5.4

Areas of future research

Orthopedic research that focuses on treatments of injuries to the bones and joints of the foot are often focused on alternatives to arthrodesis. Arthrodesis is recommended as an end stage treatment for arthritis in many locations in the foot [59]. It is also used in cases of midfoot injuries such as those in the tarsometatarsal joint. While arthrodesis can prevent further arthritic pain and provide an initial improved quality of life, the loss of mobility in the fused joint impacts the patient’s ability to walk and can lead to degradation of other joints in the foot [60,61]. This has motivated a large body of work researching alternative ways to treat arthritis and other injured/diseased states without permanent fixation. One approach has been to use orthoses (braces) to perform a “simulated arthrodesis.” While these noninvasive devices could be an alternative to surgery, direct comparison between the orthosis and the arthrodesis surgery have not been done. Another alternative to arthrodesis is arthroplasty. For end stage OA in the ankle, studies have suggested the use of total ankle arthroplasty (TAA) over arthrodesis as it seeks to maintain motion of the joint thus minimizing the risk of damage to other joints [61]. A challenge with TAA is that the procedure is more involved and complex than arthrodesis and unfortunately the current arthroplasty designs do not demonstrate long term success. Since TAA patients have more revisions than those treated with arthrodesis [61,62], additional research needs to be done to determine if the added complexity and possible complications of TAA can be overcome to provide patients with greater mobility. Despite the potential complications, TAA is a better treatment choice than arthrodesis when the ipsilateral foot/ankle has previously undergone arthrodesis, as nonunion rates of up to 38% have been reported in the second ankle [63]. Another area of research has been to find relationships between the intrinsic anatomy of the bones and the biomechanics of the foot and ankle for different injuries and treatments. This work takes advantage of improvements in computational methods to gain clinically relevant biomechanical information on the foot. The most advanced finite element models and rigid body models allow joints to be defined by their anatomy instead of simplified hinges or other joints [64]. These models have contributed to understanding the changes in stress and joint contact forces after an injury [65]. They have been used to evaluate the injured foot as well as to predict how different treatments change the biomechanics of the foot as a whole [66]. The groundwork so far in this area has pointed to several areas for model improvement. Some of these are to better characterize the material properties of healthy and diseased tissue, improve estimates of the loading conditions to the muscles for specific activities, and incorporate patient-specific anatomy and properties for the wide range of tissues in the foot [64].

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[11] MacNiel JA, Boyd SK. Accuracy of high-resolution peripheral quantitative computed tomography for measurements of bone quality. Med Eng Phys 2007;29:1096 105. Available from: https://doi.org/10.1016/j.medengphy.2006.11.002. [12] Sievanen H, Oja P, Vuori I. Precision of dual-energy x-ray absorptiometry in determining bone mineral density and content of various skeletal sites. J Nucl Med 1992;33(6):1137 42. [13] Cann CE. Quantitative CT for determination of bone mineral density: a review. Radiology 1988;166(2):509 22. Available from: https://doi. org/10.1148/radiology.166.2.3275985. [14] Khoo B, Brown K, Cann C, Zhu K, Henzell S, Low V, et al. Comparison of QCT-derived and DXA-derived areal bone mineral density and T scores. Osteoporos Int 2009;20:1539 45. Available from: https://doi.org/10.1007/s00198-008-0820-y. [15] Panchbhavi VK, Boutris N, Patel K, Molina D, Andersen CR. CT density analysis of the medial cuneiform. Foot Ankle Int 2013;34 (11):1596 9. Available from: https://doi.org/10.1177/1071100713499904. [16] Pelt C, Turner C, Bachus K, Foreman B, Beals T. Micro-CT analysis of the medial wall of the human medial cuneiform. Orthopedics 2011;34 (5):88 92. Available from: https://doi.org/10.3928/01477447-20110317-07. [17] Yoganandan N, Chirva S, Voo L, DeVogel N, Pintar FA. Foot-ankle complex injury risk curves using calcaneus bone mineral density data. J Mech Behav Biomed Mater 2017;72:246 51. Available from: https://doi.org/10.1016/j.jmbbm.2017.05.010. [18] Ketcham RA, Ryan TM. Quantification and visualization of anisotropy in trabecular bone. J Microscopy 2004;213(2):158 71. Available from: https://doi.org/10.1111/j.1365-2818.2004.01277.x. [19] Giddings V, Beaupre G, Whalen R, Carter D. Calcaneal loading during walking and running. Med Sci Sports Exerc 2000;32(3):627 34. Available from: https://doi.org/10.1097/00005768-200003000-00012. [20] Griffin N, D’Aouˆt K, Ryan T, Richmond B, Ketcham R, Postnov A. Comparative forefoot trabecular bone architecture in extant hominids. J Hum Evol 2010;59:202 13. Available from: https://doi.org/10.1016/j.jhevol.2010.06.006. [21] Tsegai Z, Skinner M, Gee A, Pahr D, Treece G, Hublin J-J, et al. Trabecular and cortical bone structure of the talus and distal tibia in Pan and Homo. Phys Anthropology 2017;163:784 805. Available from: https://doi.org/10.1002/ajpa.23249. [22] Herbert D, Lebrun R, Marivaux L. Comparative three-dimensional structure of the trabecular bone in the talus of primates and its relationship to ankle joint loads generated during locomation. Anat Rec 2012;295:2069 88. Available from: https://doi.org/10.1002/ar.22608. [23] Turunen M, Prantner V, Jurvelin J, Kroger K, Isaksson H. Composition and microarchitecture of human trabecular bone change with age and differ between anatomical locations. Bone 2013;54:118 25. Available from: https://doi.org/10.1016/j.bone.2013.01.045. [24] Mittra E, Rubin C, Gruber B, Qin Y-X. Evaluation of trabecular mechanical and microstructural properties in human calcaneal bone of advanced age using mechanical testing, micro-CT, and DXA. J Biomech 2008;41(2):368 75. Available from: https://doi.org/10.1016/j. jbiomech.2007.09.003. [25] Metcalf L, Dall’Ara E, Paggiosi M, Rochester J, Vilayphiou N, Kemp G, et al. Validation of calcaneus trabecular microstructure measurements by HR-pQCT. Bone 2018;106:69 77. Available from: https://doi.org/10.1016/j.bone.2017.09.013. [26] Gunaratne I, Singh P, Dega R. Transient migratory osteoporosis of the foot. Foot Ankle Surg 2007;13:51 4. Available from: https://doi.org/ 10.1016/j.fas.2006.07.001. [27] Calvo E, Alvarez L, Fernandez-Yruegas D, Vallejo C. Transient osteoporosis of the foot: Bone marrow edema in 4 cases studied with MRI. Acta Orthop Scand 1997;68(6):577 80. Available from: https://doi.org/10.3109/17453679708999030. [28] Mow VC, Proctor CS, Kelly MA. Biomechanics of articular cartilage. Basic biomechanics of the locomotion system. Philadelphia, PA: Lea and Febiger; 1989, p. 31 58. [29] Fox A, Bedi A, Rodeo S. The basic science of articular cartilage: Structure, composition, and function. Sports Health 2009;1(6):461 8. Available from: https://doi.org/10.1177/1941738109350438. [30] Buckwalter JA, Mow VC, Ratcliffe A. Restoration of injured or degenerated articular cartilage. J Am Acad Orthop Surg 1994;2:192 201. [31] Mow V, Ratcliffe A. Structure and function of articular cartilage. Basic orthopaedic biomechanics. Philadelphia, PA: Lippincott-Raven Publishers; 1997, p. 113 64. [32] Torzilli P, Mow V. On the fundamental fluid transport mechanisms through normal and pathological articular cartilage during function-I the formulation. J Biomech 1976;9(8):541 52. Available from: https://doi.org/10.1016/0021-9290(76)90071-3. [33] Mow V, Mansour J. The nonlinear interaction between cartilage deformation and interstitial fluid flow. J Biomech 1977;10:31 9. Available from: https://doi.org/10.1016/0021-9290(77)90027-6. [34] Mow VC, Kuei SC, Lai WM, Armstrong CG. Biphasic creep and stress relaxation of articular cartilage in compression: theory and experiments. J Biomech Eng 1980;102:73 84. Available from: https://doi.org/10.1115/1.3138202. [35] Soulhat J, Buschmann MD, Shirazi-Adl A. A Fibril-network-reinforced biphasic model of cartilage in unconfined compression. J Biomech Eng 1999;121:340 7. Available from: https://doi.org/10.1115/1.2798330. [36] Lai WM, Hou JS, Mow VC. A triphasic theory for the swelling and deformation behavior of articular cartilage. J Biomech Eng 1991;113:245 58. Available from: https://doi.org/10.1115/1.2894880. [37] Mak AF. The apparent viscoelastic behavior of articular cartilage—the contributions from the intrinsic matrix viscoelasticity and interstitial fluid flows. J Biomech Eng 1986;108:123 30. Available from: https://doi.org/10.1115/1.3138591. [38] Huang C-Y, Soltz M, Kopacz M, Mow V, Ateshian G. Experimental verification of the roles of intrinsic matrix viscoelasticity and tensioncompression nonlinearity in the biphasic response of cartilage. J Biomech Eng 2003;125:84 93. Available from: https://doi.org/10.1115/ 1.1531656. [39] Caillet R. Structural functional anatomy. Foot and Ankle Pain. Philadelphia, PA: F. A. Davis Company; 1997, p. 1 46.

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[40] Sugimoto K, Takakura Y, Tohno Y, Kumai T, Kawate K, Kadono K. Cartilage thickness of the talar dome. Arthroscopy 2005;21(4):401 4. Available from: https://doi.org/10.1016/j.arthro.2004.12.005. [41] Al-Ali D, Graichen H, Faber S, Englimeier K-H, Reiser M, Eckstein F. Quantitative cartilage imaging of the human hind foot: precision and inter-subject variability. J Orthop Res 2002;20:249 56. Available from: https://doi.org/10.1016/S0736-0266(01)00098-5. [42] Kuettner KE, Cole AA. Cartilage degeneration in different human joints. Osteoarthr Cartil 2005;13(2):93 103. Available from: https://doi.org/ 10.1016/j.joca.2004.11.006. [43] Valderrabano V, Horisberger M, Russell I, Dougall H, Hintermann B. Etiology of ankle osteoarthritis. Clin Orthop Relat Res 2009;468 (7):1800 6. Available from: https://doi.org/10.1007/s11999-008-0543-6. [44] Teng A, Pinzur M, Lomasney L, Mahoney L, Havey R. Functional outcome following anatomic restoration of tarsal-metatarsal fracture dislocation. Foot Ankle Int 2002;23(10):922 6. Available from: https://doi.org/10.1177/107110070202301006. [45] Coester L, Saltzman C, Leupold J, Pontarelli W. Long-term results following ankle arthrodesis for post-traumatic arthritis. J Bone Jt Surg 2001;83-A(2):219 28. Available from: https://doi.org/10.2106/00004623-200102000-00009. [46] Lundberg A, Svensson O, Nemeth G, Selvik G. The axis of rotation of the ankle joint. J Bone Jt Surg 1989;71(1):94 9. Available from: https:// doi.org/10.1302/0301-620X.71B1.2915016. [47] Nichols JA, Roach KE, Fiorentino NM, Anderson AE. Subject-specific axes of the rotation based on talar morphology do not improve predictions and subtalar joint kinematics. Ann Biomed Eng 2017;45(9):2109 21. Available from: https://doi.org/10.1007/s10439-017-1874-9. [48] Smith R, Reischl S. Treatment of ankle sprains in young athletes. Am J Sports Med 1986;14(6):465 71. Available from: https://doi.org/ 10.1177/036354658601400606. [49] Parr WC, Chatterjee HJ, Soligo C. Calculating the axes of rotation for the subtalar and talocrural joints using 3D bone reconstruction. J Biomech 2012;45:1103 7. Available from: https://doi.org/10.1016/j.jbiomech.2012.01.011. [50] Leardini A, Stagni R, O’Connor J. Mobility of the subtalar joint in the intact ankle complex. J Biomech 2001;34:805 9. Available from: https://doi.org/10.1016/S0021-9290(01)00031-8. [51] BodyParts3D. Japan: Life Science Integrated Database Center; 2008. [52] Sammarco VJ. The talonavicular and calcaneocuboid joints: anatomy, biomechanics, and clinical management of the transverse tarsal joint. Foot Ankle Clin 2004;9:127 45. Available from: https://doi.org/10.1016/S1083-7515(03)00152-9. [53] Melao L, Canella C, Weber M, Negrao P, Trudell D, Resnick D. Ligaments of the transvers tarsal joint complex: MRI-anatomic correlation in cadavers. Am J Roentgenol 2009;193(3):662 71. Available from: https://doi.org/10.2214/AJR.08.2084. [54] Novacheck T. The biomechanics of running. Gait Posture 1998;7(1):77 95. Available from: https://doi.org/10.1016/S0966-6362(97)00038-6. [55] Deland J, de Asla R, Sung I-H, Ernberg L, Potter H. Posterior tibial tendon insufficiency: which ligaments are involved. Foot Ankle Int 2005;26(3):427 35. Available from: https://doi.org/10.1177/107110070502600601. [56] Benirschke SK, Meinberg E, Anderson SA, Jones CB, Cole PA. Fractures and dislocations of the midfoot: Lisfranc and Chopart injuries. J Bone Jt Surg 2012;94(14):1326 37. Available from: https://doi.org/10.1053/j.jfas.2010.08.005. [57] Desmond EA, Chou LB. Current concepts review: Lisfranc injuries. Foot Ankle Int 2006;27(8):653 60. Available from: https://doi.org/ 10.1177/107110070602700819. [58] Lakin RC, DeGnore LT, Pienkowski D. Contact mechanics of normal tarsometatarsal joints. J Bone Jt Surg 2001;83-A(4):520 8. Available from: https://doi.org/10.2106/00004623-200104000-00006. [59] Daniels TR, Younger AS, Penner M, Wing K, Dryden PJ, Wong H, et al. 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Chapter 6

Muscles and Tendons Karen J. Mickle School of Environmental and Life Sciences, University of Newcastle, NSW, Australia

Abstract This chapter is mainly focused on the function of the muscles and tendons of the foot and ankle in the normal, pathological, and aging foot. The roles that the foot and ankle muscle and tendons play in the primary function of foot stability, balance, and locomotion are described. This chapter highlights the important role of the foot muscles in daily living activities, and how these structures can be compromised by aging, injury, and disease. However, we show evidence that foot muscles appear to respond quickly to increased mechanical stimuli; therefore, interventions such as exercises and footwear that target the neuromuscular training of these muscles may contribute to effective management of individuals with muscle atrophy and dysfunction.

6.1

Introduction

Many of the muscles that move the foot and the toes originate on the surface of the tibia and fibula and can be organized into the anterior, posterior, and lateral compartments. The anterior compartment comprises the tibialis anterior (TA), peroneus tertius (PT), and the extensor hallucis (EHL) and digitorum (EDL) longus muscles. These muscles all dorsiflex the ankle, and either assist with inversion (TA) or eversion (PT) of the foot, or extension of the joints of the big toe (EHL) and toes two to five (EDL). The lateral compartment comprises the peroneus brevis (PB) and longus (PL) muscles. These muscles plantarflex the ankle and evert the foot while stabilizing the lateral ankle and longitudinal arch of the foot. The most powerful movement of the ankle is produced by the posterior compartment, which plantarflexes the ankle. The triceps surae muscle group (gastrocnemius and soleus) inserts onto the calcaneus via the Achilles tendon (the largest tendon in the body) and can be assisted in plantarflexion by the much smaller (if present) plantaris muscle. Based on muscle mass, the gastrocnemius produces more force than most other muscles of the leg [1]. Deeper to these muscles are the tibialis posterior (TP) that also adducts and inverts the foot and the flexor hallucis (FHL) and digitorum longus (FDL) muscles that flex the joints of the big toe and toes two to five, respectively. Their function is to control the flexion of distal phalanges of the hallux or lesser toes, as well as plantarflexing the ankle. The intrinsic muscles that move the toes and stabilize the arch originate on the bones of the tarsus and forefoot. The muscles on the plantar aspect appear in four layers. The first layer is the most superficial to the sole and is located immediately underneath the plantar fascia. The three muscles in this layer include the following: flexor digitorium brevis (FDB), which flexes the proximal phalanges, and the abductor hallucis (ABH) and abductor digiti minimi (ABDM), which abduct the big toe and toe five respectively. The second layer contains the quadratus plantae (QP), which flexes the joints of toes two to five and adjusts the oblique pull of the FDL, and the four lumbricals that flex the metatarsophalangeal joints and extend the joints between the proximal, middle, and distal phalanges of toes two to five. The third layer contains three muscles. Flexor hallicus brevis (FHB) and flexor digiti minimi flex the metatarsophalangeal joints of the big and little toe respectively and the adductor hallucis (ADH) adducts the big toe. The three plantar and four dorsal interossei comprise the fourth and final plantar muscle layer. The plantar interossei have a unipennate morphology, while the dorsal interossei are bipennate and function to adduct/abduct the toes respectively. On the dorsal aspect of the foot there are only two intrinsic muscles—the extensor digitorum brevis (EDB) and the extensor hallucis brevis (EHB). Along with the extrinsic toe extensors, they play a role in aiding the extension of the medial four toes at the metatarsophalangeal and interphalangeal joints and extending the great toe at the metatarsophalangeal joint, respectively. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00039-1 © 2023 Elsevier Inc. All rights reserved.

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It is the attachment of these muscles to the bone via strong tendons that allows them to produce the desired movement. As the largest and strongest tendon in the human body, the Achilles tendon is the most important tendon for walking, running, and jumping. It attaches the triceps surae muscles to the calcaneus to produce plantarflexion at the ankle (e.g., raising up onto our toes). The posterior tibial tendon attaches the TP to the underside of the foot, passing behind the medial malleolus. Although it attaches a smaller muscle of the ankle and foot, this tendon is important in supporting the arch of the foot and helps turn the foot inward (inversion) during walking. On the other side of the ankle, the PB and PL tendons are an important lateral stabilizer for the ankle. The toes have tendons that flex or bend the toes down (attached on the bottom of the toes) and extend or straighten the toes (attached on the top of the toes). Only the flexor and extensor longus muscles have tendon attachments reaching all the way out to the most distal phalange, and therefore are able to flex or extend all of the toe joints. Assessing the specific function of intrinsic foot muscles is challenging due to their deep location and small size. However, the ability to measure muscle and tendon properties has improved greatly over recent years with advancements in the quality and accessibility of imaging modalities such as magnetic resonance imaging (MRI) and ultrasound. Ultrasound has been shown to be an inexpensive, reliable method for measuring muscle volume, muscle thickness, and cross-sectional area (CSA), as well as for quantifying plantar fascia thickness, especially in those with plantar fasciitis. Furthermore, the anatomical and physiological CSA of the FDB are significantly correlated with toe flexor strength of the lesser toes of both men and women [2,3]. This chapter will focus on the function of the muscles and tendons of the foot and ankle in the normal, pathological, and aging foot.

6.2

Biomechanical function

6.2.1 Normal foot As a group, the foot muscles play an important role in supporting the medial longitudinal arch and are imperative for efficient performance of activities of daily living including postural control in standing and walking. The activation of the toe flexor muscles is required for the push-off phase of human locomotion, as the heel leaves the ground and dorsiflexion of the metatarsophalangeal joint increases [4]. Toe flexor muscle strength therefore is an important factor for determining maximum walking speed and a correlation between strength of the toe flexor muscles and maximum walking speed has been observed [2]. From a sporting perspective, increased strength of the toe flexor muscles enhances athletic performance [5]. The role of the foot and ankle muscles and tendons in their primary function of foot stability, balance, and locomotion are described in more detail in the following sections.

6.2.1.1 Foot stability Historically, the plantar fascia has been considered as the most important soft tissue in supporting the longitudinal arch of the foot. This flat band of connective tissue runs along the sole of the foot with attachments from the medial tubercle of the calcaneus to the proximal phalanges. This structure functions as a major contributor to supporting the longitudinal arches of the foot by acting as a truss, where it undergoes tension when the foot bears weight. However, there is now a large body of evidence to show that even in quiet stance, the intrinsic foot musculature has a functional role in maintaining the medial longitudinal arch. For example, the ABH, which is the most medial muscle within the superficial layer of the intrinsic foot muscles, has a key role in supporting the arch. Originating on the posteromedial surface of the calcaneus, it inserts onto the medial sesamoid of the hallux and/or the proximal phalanx. Therefore, it acts as a dynamic elevator of the arch as it simultaneously produces flexion and inversion of the first metatarsal, inversion of the calcaneus, and external rotation of the tibia [6]. Due to the superficial location of this muscle, its function has been investigated more than most of the smaller and deeper muscles. The electromyographic (EMG) activity of the ABH has been shown to drop by 26.8% ( 6 13%) and resulted in a 3 mm navicular drop after a nerve block to the posterior tibial nerve [7]. Conversely, intramuscular electrical stimulation of the ABH, FDB, and QP counteracts medial longitudinal arch compression under load [8]. Further evidence of the role of ABH in arch support has been shown by disrupting the function of the muscle through fatigue, resulting in changes in navicular drop of up to 5 mm with a mean drop of 1.8 6 1.3 mm [9]. In addition to the intrinsic foot muscles, the FHL and FDL further contribute to the maintenance of the medial longitudinal arch [10] by resisting midfoot dorsiflexion associated with foot eversion and coupling with intrinsic muscle and plantar fascia function since all these structures insert into the digits [11,12].

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6.2.1.2 Balance The process of balance relies on integrating information from the visual, vestibular, and proprioceptive systems to elicit appropriate muscular responses to produce postural adjustments so that the body’s center of gravity is maintained over its base of support [13]. The foot and ankle muscles are very important in balance and postural control. In a regular side-byside stance position, balance in the antero/posterior direction is controlled by means of an ankle strategy (activity of the plantar/dorsiflexors), however as stance narrows (e.g., tandem stance) sway in the medial/lateral direction is dominated by an ankle strategy (activity of the ankle invertors/evertors) [14]. In terms of specific foot and ankle muscles, the ABH appears to play a particularly important role, whereby the size of the muscle was correlated to sway [15]. A larger ABH benefited open-loop dynamic stability, as well as supported a more efficient transfer from open-loop to closed loop control mechanisms. Interestingly, they also found that a larger peroneus longus/brevis muscle was correlated to increased movement of the COP during single-leg standing. Although another study found that fatiguing the dorsiflexor muscles, that is, TA and peroneal muscles causes an increase in postural sway during single-leg standing [16]. Individuals with chronic ankle instability (CAI), which could be considered as a balance issue, had smaller total extrinsic foot and ankle muscle volumes compared with age- and sex-matched healthy controls (CAI, 9.62 6 0.39 cm3/m kg; healthy, 11.13 6 1.33 cm3/m kg) [17]. The superficial posterior compartment appeared to be the most affected extrinsic muscle group (CAI, 5.15 6 0.55 cm3/m kg; healthy, 6.21 6 0.73 cm3/m kg) with the soleus muscle significantly smaller (CAI, 2.62 6 0.30 cm3/m kg; healthy, 3.26 6 0.54 cm3/m kg). Individuals who have suffered a lateral ankle sprain had a smaller CSA of the PL [18]. The CAI group also presented with smaller muscle volumes for the oblique head of ADH (CAI, 0.07 6 0.01 cm3/m kg; healthy, 0.13 6 0.03 cm3/m kg) and FHB (CAI, 0.06 6 0.02 cm3/m kg; healthy, 0.12 6 0.02 cm3/m kg) [17]. Atrophy of the FHB and oblique head of ADH in this small cohort is in agreement with the previous studies and reinforces the importance of the first metatarsal during balance and other functional tasks. The toes are ideally placed to help assist the management of the ankle balance strategy as they are the furthest away from the ankle center; therefore, load under the toes has a large ankle moment arm. Hence, intramuscular EMG activity recorded from ABH, FDB, and QP has shown that activation of these intrinsic foot muscles increases with increasing postural demand (i.e., single leg stance compared to double leg stance; Fig. 6.1) [19]. This suggests that the intrinsic foot muscles are recruited in a highly co-ordinated manner to stabilize the foot and maintain balance in the medial/lateral direction, particularly during single leg stance. These studies suggest that the morphology of foot muscles plays an important role in balance performance, and that strengthening the intrinsic foot muscles may be an effective way to improve balance. A foot strengthening program for older people found a significant increase in single-leg balance time with eyes open (15 seconds) and closed (13 seconds) [20].

6.2.1.3 Locomotion The plantar fascia also has an important dynamic function during gait. During the support phase of walking, the plantar fascia continuously elongates (by up to 12%) [21] and behaves like a spring, which may assist in conserving energy. When the toes are dorsiflexed in the propulsive phase of gait, the plantar fascia becomes tense, and significantly thinner [22], resulting in elevation of the longitudinal arch and shortening of the foot. Given that many of the intrinsic foot muscles run parallel with the plantar fascia along the longitudinal arch, it is reasonable to suggest that they play a similarly important role in dynamic locomotion activities. It has been traditionally thought that the plantar fascia contributes the most to the elastic storage and return of energy of the longitudinal arch during gait. However, there is recent evidence that suggests that the intrinsic foot muscles have the potential to actively modulate the energetic function of the foot. Kelly et al. found that the FDB muscle fascicles contract in an isometric manner when under load (equivalent to those experienced during walking), facilitating elastic energy storage in the tendon of the muscle, in addition to the energy stored within the plantar fascia [23]. In addition, ABH, FDB, and QP actively lengthen and shorten during treadmill walking and running [24]. Electrical stimulation of the intrinsic foot muscles has shown that these muscles have the capacity to support the medial longitudinal arch by slowing the deformation of the arch during the stance phase of gait [25]. In terms of the extrinsic muscles, TA eccentrically contracts in early stance to allow gradual plantarflexion of the foot and to decelerate downward motion of the foot. It then has to concentrically dorsiflex the ankle during swing for foot clearance and placement. FHL also plays an important role in the coupling between the hindfoot and first metatarsophalangeal joint kinematics [26]. The muscle tendon unit length and peak EMG of ABH, FDB, and QP increases significantly with increasing gait velocity (Kelly et al. [24], Fig. 6.2). This suggests that they are capable of contributing to the stiffness of the longitudinal arch during locomotion. EMG of these muscles showed evidence of activation during late swing phase and stance,

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FIGURE 6.1 Comparison of the electromyographic (EMG) activity for the abductor hallucis (AH or ABH elsewhere in the chapter), flexor digitorium brevis (FDB), and quadratus plantae (QP) muscles during single leg standing (SLS) and double leg standing (DLS). Adapted from Kelly LA, et al. Recruitment of the plantar intrinsic foot muscles with increasing postural demand. Clin Biomech 2012;27(1):46 51.

which may be an important mechanism to stiffen the longitudinal arch in preparation for high deformation forces associated with running. When running, the FDB is primarily active in stance phase with a large burst of activity commencing at foot contact, with a midstance peak and deactivation shortly after toe-off [27]. These findings suggest that the intrinsic foot muscles may activate during the stance phase of running to increase energy absorption and return, while also preventing excessive deformation of the midfoot when acutely transitioning to a forefoot technique.

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FIGURE 6.2 Comparison of the electromyographic (EMG) activity for the abductor hallucis (AH or ABH elsewhere in the chapter), flexor digitorium brevis (FDB), and quadratus plantae (QP), and change in muscle-tendon unit length during walking and running. FC, Foot contact; TO, toe off. Adapted from Kelly LA, Lichtwark G, Cresswell AG. Active regulation of longitudinal arch compression and recoil during walking and running. J R Soc Interface 2015;12(102):20141076.

The position of the foot at initial contact influences the capacity of foot muscles to produce torque between bone segments of the foot [27]. Although a large body of evidence exists on the biomechanical differences between runners with different foot strike patterns, whether adopting one or the other foot strike pattern may prevent or enhance longterm anatomical foot adaptations is still unknown. The majority of runners adopt a rearfoot strike (RFS) pattern landing on the heel [28]. (Editors note: Elsewhere in this book, we use “hindfoot” exclusively, but due to overwhelming convention in the existing literature, we will use “rearfoot” strike in this chapter.) Due to the construction of traditional cushioned footwear, a RFS runner is likely to develop a reliance upon the extrinsic mechanical properties that the shoe mid-sole provides. In contrast, runners with forefoot strike (FFS) patterns (i.e., those who tend to land on the ball of the foot) rely less on the shoe properties [29] and utilize foot biological properties to control impact forces [30]. Therefore, the magnitude and location of the external ground reaction force may change the recruitment of muscles around a joint [31]. Proponents of barefoot (and, as a consequence, forefoot running) have hypothesized that this strike pattern is likely to recruit intrinsic foot muscles to a greater extent than their rearfoot striking counterparts. When controlling foot strike pattern in non-FFS runners, changing to a FFS pattern results in increased intrinsic foot muscle activation of ABH and FDB foot muscles compared to a RFS pattern. While running barefoot on a treadmill ABH shows a large burst of activity, commencing during late swing phase and continues through to toe-off. The peak activation coincides within midstance, as the foot absorbs the load. When running with an FFS pattern, the mean ABH activity is greater in swing and stance phases, and the mean stance-phase activation for FDB is higher, than when running with a RFS pattern in runners who habitually strike the ground with a heel-strike pattern [27]. However, the peak activations were not significantly different. This study showed the immediate effects of changing foot strike pattern on muscle activation, however, could not address the longer-term adaptation effects of foot strike pattern on muscle properties of the foot. Sprinters (known to adopt an FFS pattern) have more developed foot muscles. Ultrasonography-measured thicknesses in the leg and foot muscles are larger in sprinters than in non-sprinters (Fig. 6.3), however the size is not positively related to greater sprint performance, in fact, a thicker ABH was associated with a slower 100 m sprint time [32]. At the other end of the running spectrum, there are millions more people engaging in long-distance running. To help distinguish whether foot muscles develop because of running, or by the manner of runner style, we compared RFS long distance runners to FFS long distance runners. Using ultrasound imaging and a toe flexor strength dynamometer, we quantified differences in foot muscle size (CSA and thickness) and toe flexor strength in 20 male, experienced, recreational long distance runners. Participants were classified as having a RFS (n 5 10) or FFS (n 5 10) pattern based on their habitual foot strike pattern tested on an instrumented treadmill; running at preferred speed and wearing their habitual running shoes. Interestingly, we found no differences in muscle or tendon size (Fig. 6.3), or strength of the toe flexors (FFS 5 0.019 6 0.008 vs RFS 5 0.021 6 0.008 BW/m) [33], indicating that foot strike does not affect the size of foot muscles and their ability to produce force.

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FIGURE 6.3 Ultrasonography-measured thicknesses in the leg and foot muscles are larger in sprinters than in non-sprinters, with no difference between rearfoot strikers (RFS) vs forefoot strikers (FFS) long distance runners. * Indicates significant difference between sprinters and controls. ABH, Abductor hallucis; FDB, flexor digitorum brevis; FDL, flexor digitorum longus; FHB, flexor hallucis brevis; FHL, flexor hallucis longus; GL, lateral gastrocnemius; GM, medial gastrocnemius; PLB, peroneus longus and brevis; TA, tibialis anterior.

Similarity in the size of the intrinsic foot muscles between the two foot strike groups in our study is in agreement with recent research from Farris and colleagues [34] who suggested that the intrinsic foot muscles have minimal effect on the stiffness of the arch when absorbing high loads. Given the experience of the runners in our study ( . 40 km/wk for the last 5 years), it is likely that the muscles were fully adapted and contribute equally to propulsion, independent of foot strike. However, it should be noted that the group with a FFS pattern runs habitually in traditional or less supportive shoes, but not minimal shoes or barefoot. The similarity with RFS may partially depend on shoe assistance. There is some evidence to suggest that the type (or lack) of footwear influences the size of the foot muscles (see Footwear and orthoses section below).

6.2.2 Aging Many aspects of declining physical functioning, including muscle strength, are regarded as an inevitable consequence of aging, with age-related loss of muscle mass one the main determinants of frailty. Muscle atrophy in older adults has been detected in numerous muscles of the lower limb including the triceps surae muscles [35,36], with the decline in muscle strength appearing around the sixth decade. Loss of muscle mass in older adults has been associated with an annual decline in lower leg strength of approximately 3% [37]. Muscles within the feet, including those that control the toes, also suffer from atrophy and an associated loss of strength with aging. Both the size and strength of the toe flexor muscles are significantly reduced in older adults (67.1 6 2.9 years) compared to younger counterparts (28.8 6 8.2 years) [3]. The thickness and CSA of FHB, QP, ABDM, FDL, FHL, and FDB muscles were reduced by 19% 45% in the older participants (Fig. 6.4). This was accompanied by a 38% and 35% reduction in strength of the hallux and lesser toes, respectively. Our finding was similar to the 29% lower toe strength displayed by older people when performing a maximal reach task [38] and the 27% 32% reduction in toe strength in older people reported by Menz et al. [39]. The gastrocnemius muscles were up to 15% thinner in women aged 60 years or older compared to their younger counterparts, and muscle atrophy was detected in men aged 50 or older [35].

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FIGURE 6.4 Cross-sectional area (CSA) of foot and leg muscles of young and older adults. * Indicates significant difference between groups. ABDM, Abductor digiti minimi; ABH, abductor hallucis; FDB, flexor digitorum brevis, FDL, flexor digitorum longus, FHB, flexor hallucis brevis; FHL, flexor hallucis longus; QP, quadratus plantae.

This reduction in toe flexor strength can greatly affect older people’s ability to walk safely, whereby weak toe flexors increase the risk of falls in older people. The simple “paper grip test” (pass/fail test of ability to hold onto a piece of paper while assessors attempts to remove), a clinical measure of toe strength, was an independent predictor of falls in 176 retirement village residents [40]. Furthermore, our research group investigated whether reduced toe flexor strength increased the risk of falling in 312 community-dwelling older adults (aged 60 90 years) [41]. Toe flexor strength was quantified at baseline and the participants were then followed prospectively to determine their incidence of falls over 12 months. In total, 107 (35%) participants experienced a fall in the 12-month period of the study. Compared to non-fallers, the fallers had significantly less strength of the hallux and lesser toes (a reduction of more than 20%). As quadriceps and ankle strength did not differ between the fallers and non-fallers, the reduced toe strength recorded for fallers was deemed unlikely to be a marker of generalized lower limb muscle weakness. Furthermore, hallux strength is a stronger predictor of falls than other more commonly measured falls risk factors, such as age, gender, falls risk score and quadriceps strength. In fact, each unit (% body weight) increase in hallux strength decreased the odds of sustaining a fall by 7% [41].

6.2.3 Pathologies Normal function of the muscle and tendons of the foot can be affected by several pathologies. The effect of some of the most common foot pathologies, such as plantar fasciitis, pes planus, toe deformities and diabetic neuropathy, are discussed in this section.

6.2.3.1 Plantar fasciitis If excessive and repetitive tensile forces are imposed onto the plantar fascia, presumed development of microtrauma results in a condition known as plantar fasciitis. Plantar fasciitis is a common musculoskeletal disorder characterized by pain involving the inferomedial aspect of the heel that is exacerbated following periods of non-weight bearing. Diagnostic ultrasonography of the plantar fascia, typically demonstrates diffuse or localized hypoechoic (gray/black; Fig. 6.5) areas within a thickened calcaneal attachment [42]. Several studies agree that a plantar fascia thickness over 4 mm measured via ultrasound is consistent with plantar fasciitis. The involved plantar fascia was 1.2 to 1.3 mm thicker than the uninvolved side in people with unilateral plantar fasciitis [22], whereas side-to-side difference in plantar fascia thickness tends to be less than 0.1 mm in people without the condition [22].

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FIGURE 6.5 Comparison of a healthy plantar fascia (cross-sectional view) and a patient with plantar fasciitis with the hypoechoic area highlighted.

Although the causes of plantar fasciitis are clearly multifactorial, weakness of the intrinsic foot muscles is thought to play a role in its development. Individuals with unilateral plantar fasciitis have been shown to have weaker toe flexor strength than a control group, and their affected foot produces 8% less force than their non-affected foot [43]. Few studies have been able to determine whether plantar fasciitis is associated with atrophy of the important arch-supporting muscle groups. In a cohort of unilateral chronic plantar fasciitis patients, axial MR images were acquired of the plantar intrinsic and TP muscles [44]. Compared to the unaffected foot, plantar fasciitis feet exhibited a 5.2% reduction of forefoot muscle (63.4 6 14.8 cm3 vs 67.5 6 18.9 cm3). However, in the hindfoot, no significant muscle size differences were found between the unaffected (45.8 6 17.4 cm3) foot and the symptomatic foot (44.6 6 13.3 cm3). There were also no significant differences between unaffected and plantar fasciitis feet in total muscle volumes (113.3 6 18.2 cm3 and 108.0 6 14.1 cm3 respectively). Furthermore, the maximal CSA of TP did not differ between legs [44]. As this study did not have a control group without plantar fasciitis, it is difficult to draw strong conclusions about the role of muscle weakness in the development of the condition. Plantar fasciitis is a relatively common injury that occurs in running athletes. In contrast to the aforementioned study, runners with chronic plantar fasciitis have been shown to have smaller intrinsic foot muscle volumes in the hindfoot region compared to healthy runners [45]. However, similar to Chang et al. [44] hindfoot and forefoot segments were defined by splitting the total number of images containing muscle into posterior half and anterior half. Therefore, we do not know which muscles are specifically affected in the condition.

6.2.3.2 Pes planus Characteristics associated with pes planus, including a lowered medial longitudinal arch and increased dorsiflexion, eversion, and external rotation of the foot, are often thought to increase the risk of injury and lower limb musculoskeletal conditions. However, the exact mechanism by which planus foot type may lead to an elevated risk of lower limb pathology is not completely understood. Now that we have a greater understanding of the important role that the foot muscles play in supporting the medial longitudinal arch, the contribution of muscles to foot posture, and thus planus feet, has been the focus of several studies. The extrinsic foot muscles including TP, TA, FHL, and FDL provide additional support for the medial longitudinal arch [46]. Atrophy of the TP and compensatory hypertrophy in FDL (suggesting greater activity) has been noted on MRI in cases of pes planus that are associated with TP tendon insufficiency [47]. Those with flat-arched feet display altered muscle activation patterns, such as, increased amplitudes of TA (at contact) and TP (during midstance/propulsion), and reduced activity of PL and PB (in early stance) compared to individuals with normal-arched feet [48 50]. This combined activation pattern therefore favors the invertor muscles on the medial aspect of the ankle. This seems contrary to the fact that by inserting on the plantar aspect of the first metatarsal, the PL might be able to plantarflex the metatarsal and thereby elevate medial longitudinal arch height. However, its moment arm for this function is likely very small and combined with the small muscle volume compared to other leg muscles, it seems unlikely that PL contributes significantly to the arch height of the foot.

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In agreement with the MRI study by Wacker et al. [47], the CSA of FDL was 17% larger in runners with planus feet compared to runners with normal feet. No significant differences were observed in the CSA and thickness of PL and PB [51]. We investigated this further with a much larger cohort of participants with pes planus (n 5 49), imaging the FDL, FHL, PL, PB, and TA with ultrasound [52]. The planus feet had significantly larger extrinsic invertor muscles (FDL and FHL), with the CSA of FDL and FHL greater by 28% and 24%, respectively. Measured thicknesses of FDL (15%) and FHL (10%) muscles were similarly greater in the pes planus group compared to controls (n 5 49). However, the peroneal muscles were significantly smaller in the pes planus group. The CSA and thickness of the peroneal group was decreased by 14.7% and 10% respectively. The reduced peroneal muscle size in our pes planus participants is in agreement with Murley et al.’s [49] report of decreased peroneal muscle activity in flatfeet during walking. However, in a more recent study, the same group found a 7.6% association between a thicker peroneous longus muscle and a flatter foot type [53]. As a primary evertor of the hindfoot, reduced action of the peroneal muscles should advantage the inverter muscles such as TP. This suggests that the flatter foot posture disadvantages FHL and FDL so that they need to generate greater forces to contribute the required moments and thus facilitate normal sagittal plane ankle function. This in turn, may lead to hypertrophy of the FHL and FDL as supported by the three studies [47,51,52]. Thickness of TA does not appear to be associated with foot posture [51 53]. Anesthetic paralysis and deliberate fatiguing of plantar intrinsic muscles results in reduced medial longitudinal arch height [7,9], though these experimental approaches do not indicate how individual plantar structures contribute to arch integrity. Interestingly, ABH appears to be larger in planus feet[51,54]. Recreational runners classified with a planus foot type (n 5 9) have a larger ABH and FDB and smaller ABDM than runners with a normal foot-type [51]. The thickness of the ABH was larger by 7.5%, and the CSA of FDB 18.7% larger. However, the thickness and CSA of ABDM were smaller by 10% 12%, in individuals with planus feet. Conversely, our larger assessment of 49 individuals with pes planus found the CSA of ABH and FHB muscles to be significantly smaller (ABH 212.8%, FHB 28.9%) compared to the normal group. Thickness of these muscles was likewise smaller (6.8%, 7.6% respectively) [52]. In a more recent study on older people, a smaller FHB was associated with reduced navicular height in standing and a greater navicular drop [55]. No association between size of FHB and navicular height was present while in the seated position, so this again suggests that intrinsic foot muscles play a more important role in supporting, or stiffening, the medial longitudinal arch under loaded conditions. Compensations for fatigue might explain these observations of smaller intrinsic foot muscles in people with pes planus. Muscle fatigue decreases power output and reduces work capacity. Premature and increased activity of ABH and FHB during stance phase in pes planus was demonstrated by Mann and Inman [4], perhaps suggesting increased risk of fatigue as when muscle fibers are repeatedly activated. Therefore, fatigue in the plantar intrinsic muscles could result in a loss of support for the medial longitudinal arch and result in compensatory action from FDL and FHL. Electrical stimulation of the ABH during gait has been shown to slow the deformation of the medial longitudinal arch in flat-footed subjects [56], which does suggest that this muscle may be less efficient in the lower arch, but has the capacity to respond. In support of this, it appears that we can increase the CSA and strength of ABH and FHB after strengthening exercises combined with foot orthoses [57]. In terms of the tendinous structures of the foot and ankle, an association between a flatter foot types and thicker TA tendon has been reported, whereby morphometry accounted for 7% of the variation in foot type [53]. However, they found the inverse association with the thickness of the Achilles tendon, which explained 16% of variation, whereby flatter foot types had a thinner Achilles. Although we found no difference in the calcaneal portion of the PF, the middle (210.6%) and anterior (221.7%) sections of the plantar fascia were thinner in planus feet compared to controls [52]. Furthermore, dysfunction of the posterior tibial tendon (e.g., tendinopathy) is often partially implicated in cases of adult acquired flat foot. As the tendon loses its ability to maintain the integrity of the medial longitudinal arch in conjunction with other ligaments and joint capsules gradually becoming weaker, a flatfoot develops.

6.2.3.3 Toe deformities Toe deformities are highly prevalent among older people with up to 60% 74% of older adults having lesser toe or hallux valgus deformities [58,59]. Atrophy of the plantar foot muscles, and the associated development of an imbalance between the flexor and extensor muscles, is believed to be a primary cause of toe deformities such as claw and hammer toes, and prominent metatarsals [60,61]. It is thought that weak toe flexor muscles may not be able to counterbalance the toe extensors, which will lead to phalangeal joints maintaining an extended resting position, creating the deformity. Historically, the anatomy of these toe deformities had been investigated by cadaveric studies such as those by Myerson and Shereff [62]. They reported that when the metatarsophalangeal joints were hyperextended, in the case of claw and hammer toes, this caused the toe extensor tendons to shift dorsally and form “bowstrings” (Fig. 6.6). They also found that the interossei were subluxated dorsal to the transverse axis of the metatarsal head by up to 8 mm in claw and hammer toes. The

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FIGURE 6.6 Attachment of the EDL, EHL, FDL, FDB, ABH, and FHL tendons in a healthy foot (A, C) and altered muscle/tendon pull on toes with clawing (B) and hallux valgus (D). ABH, abductor hallucis; EDL, extensor digitorum longus; EHL, extensor hallucis longus; FDB, flexor digitorum brevis; FDL, flexor digitorum longus; FHL, flexor hallucis longus.

orientation of the lumbrical tendons were also shifted by up to 50 degrees with respect to the metatarsal shaft. This study highlighted the spectrum of these toe deformities, whereby there is progressive ranges of subluxation of the proximal phalanges and associated adaptive shortening of the skin and articular soft tissues. This joint malalignment will alter the axis of the intrinsic toe flexors, in turn making them less biomechanically efficient (Fig. 6.6). Locke [11] measured muscle fiber CSA of FDB in two cadavers with claw toes and found that they had a larger fiber CSA than two cadavers with neutrally aligned toes, but there were obvious limitations to this study (small sample size, alterations to tissue associated with embalming and histological processing, and lack of cadaver medical history). With modern imaging techniques, several more studies were conducted in vivo. A sonographic study showed that the thickness and CSA of the ABH muscle were reduced in people with hallux valgus [63], however this was the only foot muscle measured. Further research has shown that older females with moderate-to-severe hallux valgus have significantly smaller ABH and FHB muscles than females without any toe deformities [64]. This is in agreement with other studies that have reported that people with hallux valgus had significantly reduced hallux flexor strength [41,65], hallux abduction strength [65], and a smaller ABH CSA [63] than those without the deformity. However, the mechanism of the associations remains unknown. It is possible that in more severe hallux valgus deformity, excessive tension on the EHL tendon would hold the hallux in extension and reduce the capacity of the flexor muscles to actively contract against the ground (Fig. 6.6). Alternatively, if atrophied through disuse, ABH cannot adequately resist the pull from ADH which may lead to adduction of the phalanx (toward to midline of the foot). This is supported by the findings of a marked decrease in the muscle activity of ABH during abduction when compared with adduction of ADB in people with hallux valgus [66]. Similarly, older men and women with lesser toe deformities have significantly smaller QP, FDB, ABH, and FHB muscles than individuals without any deformities [64]. The 9% 25% reduction in size of these muscles support the findings of previous research, whereby older people with lesser toe deformities had significantly reduced strength of the lesser toes compared to those without toe deformities [41]. Interestingly, it appears that the intrinsic toe flexor muscles are atrophied, but the extrinsic FDL muscle does not differ in size between those with and without toe deformities. It is the role of the plantar intrinsic muscles to stiffen the joints of the foot and help keep the toes flat on the ground [67,68] whereas the long flexor muscles cross the ankle joint therefore play a role in plantarflexion of the ankle as well as power and control gait accelerations. Therefore, it is possible that the former is most affected in people with lesser toe deformities.

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6.2.3.4 Diabetic neuropathy Diabetic polyneuropathy (DPN) is one of the most common complications associated with diabetes, present in 50% 70% of older people with the disease [69]. The presence of neuropathy in people with diabetes has a major detrimental effect on the foot and ankle muscles, and atrophy of these muscles is correlated to the severity of neuropathy. Motor neuropathy is believed to lead to weakness in the intrinsic muscles of the foot, thus upsetting the delicate balance between flexors and extensors of the toes. Given the progression of DPN follows a distal-to-proximal pathway, deficits in strength and balance are most evident at the foot and ankle. Through the use of MRI, the total volume of the foot muscles appears to be around half that in people with diabetes compared to non-diabetic controls [70,71]. This muscle atrophy results in functional impairment whereby foot and ankle muscle strength is significantly reduced in the patients with DPN. In later stage DPN, sensory and motor dysfunction occurs and muscle weakness develops, with muscle weakness progressing with the level of neuropathy [69]. Of concern, neuropathic patients have a 41% reduction in the strength of their ankle dorsiflexor and plantarflexor muscles compared to non-neuropathic patients [72]. This loss of strength was associated with a 32% reduction in the volume of these lower leg muscles. A length dependant neuropathic process was confirmed; muscle atrophy was most pronounced at the distal lower leg (CSA reduced by 65%) compared to the mid lower leg (CSA reduced by 43%) and proximal lower leg and thigh (no change in CSA) [72]. Furthermore, total volume of the intrinsic foot muscles is halved in patients with DPN [70], confirmation of this length dependency. Longitudinal examination of the same patients has found an annual decline of 4.5% 5% for the ankle dorsiflexors and plantarflexors, a significantly faster decline than aging alone (1.7% 1.8%) [73]. Foot muscle volume declined by 3% per year in neuropathic patients compared to 0.2% in healthy adults (Fig. 6.7) [73]. Specifically, EDB in diabetic patients was 46% smaller in CSA and 29% thinner, and the thickness of the muscle group between the first and second metatarsals was 26% smaller [74]. Similarly, Bus et al. [75] reported a 73% reduction in the CSA of the forefoot muscles in diabetic patients compared to controls using MRI. Our research team has recently investigated the strength of the foot and ankle muscles in 40 community-dwelling men and women aged 50 years or greater with type 2 diabetes (68 6 8.6 years). Similar to the previous studies, the foot and ankle muscle strength were significantly reduced in the participants with DPN, whereby toe and ankle strength were up to 36% and 48% lower than those without neuropathy, respectively (unpublished).

6.2.4 Footwear and orthoses The influence of footwear on muscles and tendons of the foot has been a topic of great debate. As muscles and tendons are highly malleable tissues, habitual use of footwear or shoe inserts might induce structural and functional changes in the muscles

FIGURE 6.7 Magnetic resonance images through the proximal part of the lower leg and the midfoot region in a neuropathic diabetic patient at baseline (A, C) and approximately 10 years later (B, D). E H Corresponding images for a control participant at baseline (E, G) and at follow-up (F, H). From Andreassen CS et al. Accelerated atrophy of lower leg and foot muscles—a follow-up study of long-term diabetic polyneuropathy using magnetic resonance imaging (MRI). Diabetologia 2009;52(6):1182 91.

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and tendons due to different joint positions of the foot and ankle. Long term adoption of specific footwear (or lack thereof) and orthoses, and whether they may prevent or enhance long-term anatomical foot adaptations is still relatively unknown. While carefully controlled studies to investigate the influence of footwear on foot muscles are difficult to conduct, several researchers have compared the foot structure of people from predominately unshod populations to those in more traditional western environments. Holowka and colleagues collected data from minimally-shod men from northwestern Mexico (Tarahumara) and men from urban/suburban areas in the United States who wore “conventional” shoes [76]. Compared to the conventionally-shod men from the US, the Tarahumara men had higher and stiffer longitudinal arches, and larger ABH and ABDM muscles. The authors concluded that the use of conventional modern shoes is associated with weaker intrinsic foot muscles that may predispose individuals to reduced foot stiffness and potentially flatter feet. The often-suggested hypothesis that standard running shoes may contribute to atrophy of the intrinsic foot muscles is conjectural. There have been too few high-quality studies to determine this. This may be primarily due to the challenges of measuring the force production of these small muscles of the foot. Some studies have shown changes in morphology and function of toe flexor muscles after training in specific athletic footwear [77 79]. In a five-month prospective study, an athletic group used a barefoot-like minimal shoe during warm-up training while the control group used conventional training shoes for the same training program. The minimal shoe group showed an increase of toe flexor strength of nearly 20% and an increase of anatomical CSA of the toe flexors of 5% [79]. Use of minimalistic shoes has also been associated with a thicker ABH muscle, a thinner proximal plantar fascia, and thicker Achilles tendon compared to other shoe types [80]. These findings suggest that flexible shoes might increase mechanical stimuli on foot structures and lead to musculoskeletal adaptations. Conversely, transitioning into minimalistic footwear has shown to produce a greater risk of bone marrow edema in the metatarsals [81]. A group of healthy runners (n 5 33) that were randomized into two groups (traditional running footwear or minimal support footwear) had MRI scans of their feet before and after a 12-week training program. Following the running program, both groups increased the CSA and volume of the FDB by 11% and 21%, respectively, however the minimally shod runners also had significant increases in the CSA and volume of the ABDM by 18% and 22%, respectively, and this was accompanied by a 60% increase in longitudinal arch stiffness [82]. In terms of recreational/occupational footwear, habitual wearing of high heeled shoes has been shown to affect the properties of the foot and ankle muscles and tendons. Wearing high heels places the calf muscle tendon unit in a shortened position. According to a previous small study, women who regularly wear high heels (n 5 11) have shorter gastrocnemius medialis fibers (49.6 6 5.7 mm vs 56.0 6 7.7 mm) and thicker and stiffer Achilles tendons than women who wear flat shoes (n 5 9) [83]. To understand the structural and functional effects of wearing foot orthoses in shoes, several researchers have investigated lower limb muscle activity response to the presence of orthoses. Orthoses reduce TP activity level during gait in both normal and flatfooted participants [84,85]. Murley et al. investigated whether modified prefabricated or customized foot orthoses influence lower limb muscle activity, and if so, whether they optimize or “reverse” the abnormal lower limb muscle activity previously observed in people with flat feet [85]. The only consistent finding was that both styles of foot orthoses significantly decreased TP EMG amplitude (13% 19%) during the stance phase compared with the shoe-only condition. This may suggest that there is a decreased demand for TP when the foot is supported by foot orthoses, however, it is unknown whether this may actually be functionally beneficial in people with flat feet. Interestingly, only the prefabricated foot orthoses had a significant effect on PL EMG amplitude, with a 21% 24% increase in muscle activity during the push-off phase compared to the shoe only condition [85]. Similarly, Akuzawa and colleagues found that the muscle activity level of the PL and FDL remained the same for all conditions (barefoot, with footwear, and with orthoses) in each phase [84]. However, owing to the varying types and design of orthoses, and the lack of universally accepted guidelines for orthotic design or prescription, there is no definitive consensus on the influence of orthoses on the structure and function of the foot muscles and tendons and further research is warranted.

6.2.5 Interventions This chapter has highlighted the important role that the foot muscles play in activities of daily living and how these structures can be compromised by aging, injury, and disease. However, it has also been highlighted that foot muscles appear to respond quickly to increased mechanical stimuli. Therefore, it is plausible that interventions that target the neuromuscular training of these muscles may contribute to effective management of individuals with muscle atrophy and dysfunction. In fact, there is strong evidence that strength training can increase toe strength. Kokkonen et al. had 38 college athletes complete a toe resistance training program, using weights on a pulley system 3 days a week for 12 weeks [86].

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Compared to an age- and gender-matched control group, toe flexor strength increased by approximately 46% and resulted in a significant improvement in vertical jump performance. Similarly, Unger et al. employed a within-subject design whereby 15 men and women (aged 21 62 years) completed three training sessions of toe strengthening exercises (Archxerciser) on one foot per week for 6 weeks [87]. Compared to the “untrained” (control) foot, toe flexor strength was improved almost fourfold after 6 weeks. Importantly, these unilateral strength gains were accompanied by significant improvements in single limb horizontal and vertical jump performance (on the trained foot side). The most impressive toe strength gains have been in a study that resulted in a 60% 70% increase in toe flexor strength after 7 weeks of high intensity strength training [88]. These studies clearly confirm that strength training can increase toe flexor muscle strength and performance in younger cohorts. Other research has associated toe exercises with toe strength gains and improved motor task performance in older adults. Nine nursing home residents (men and women, aged 77.8 6 2.8 years) undertook toe-grasp training for 10 minutes, 3 times a week [89]. This training required participants to use their toes to gather a towel attached to a weight and pass beanbags from one place to another, using their toes. After 8 weeks of training, the toe-grasp training group significantly improved the participants’ spontaneous sway performance with both their eyes open and closed by 16% and 26%, respectively. In a study of healthy, older people (n 5 117), participants in the intervention group attended three group exercise classes each week for 12 weeks. During each class, an exercise physiologist led the participants through a series of 10 exercises, which were predominantly designed to strengthen the toe flexor muscles (Fig. 6.8). Compared to baseline, participants in the intervention group significantly increased their hallux and lesser toe strength (up to 36%), whereas there was no significant change in toe strength in the home or control groups (see Fig. 6.9) [20]. Additional benefits of the Toe Training program for participants in the intervention group included a significant increase in single-leg balance time with eyes open (15 seconds) and closed (13 seconds). Both foot exercise groups appeared to have small improvements in their balance after completing their respective programs. Only the Toe Training group showed a significant improvement in the balance time with their eyes open, but both groups improved the length of time they could balance on one-leg with their eyes closed. The only exercises that were replicated in both exercise programs were the short foot exercise and heel raises. These were also the only exercises that were weight bearing and progressed to being performed on a single leg, which may explain why both groups improved their single-leg balance time. Effective evidence-based interventions to increase foot strength and restore normal function are required for those individuals whereby muscle weakness leads to dysfunction such as those with diabetic neuropathy. Based on the promising interventions mentioned earlier, my research team recently conducted a pilot study to determine whether people

FIGURE 6.8 Example exercises in the toe strengthening training program.

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FIGURE 6.9 Changes in hallux and lesser toe strength following supervised, or home exercise program. * Indicates significant difference from pre to post intervention. Adapted from Mickle KJ, et al. Efficacy of a progressive resistance exercise program to increase toe flexor strength in older people. Clin Biomech 2016;40:14 9.

with DPN would respond to the same foot strengthening program. A total of 26 participants presented with peripheral neuropathy were eligible to be randomized to the foot strengthening exercise program (n 5 15) or control groups (n 5 11). After 12 weeks, 20 participants returned to the laboratory for retesting (77%). We found a significant interaction between groups and time, whereby participants in the intervention group increased their toe strength while the control group became weaker compared to baseline (unpublished). This study has provided good pilot data to show that a foot strengthening exercise program may be suitable for people with diabetic neuropathy. Exercises such as the short foot exercise (Fig. 6.8, top left) have been advocated to strengthen the intrinsic foot muscles. Indeed, a study that combined the short foot exercise with an orthotic prescription found that individuals with pes planus exhibited significant increases in hallux flexor strength (14%) and the CSA of the ABH muscle (6%) [57]. However, given the attachment locations of the ABH, exercises targeting this muscle should involve diagonal movements combining both abduction and flexion of the big toe. The toes-spread-out exercise showed significantly greater activation of the ABH than did the short foot exercise (mean difference 5 45% MVC). There was no significant difference between the two exercises in activating the ADH. This resulted in the ratio of ABH to ADH muscle activity being significantly higher in the toes-spreadout exercise than in the short foot exercise [90] and may be an beneficial exercise for people with hallux valgus. A followup study employing the toe-spread-out exercise for people with hallux valgus found a significant decrease of 3.41 degrees in the average HV angle after an 8-week intervention [91]. This was accompanied by an increase in CSA of the ABH muscle. Therefore, individualizing foot strengthening exercises for specific foot problems may be effective of preventing, treating and/or reducing the severity of these pathologies and injuries. The benefits of foot strengthening in otherwise healthy individuals require further investigation. A recent study has conducted an eight-week intervention focused on strengthening the intrinsic foot muscles in 14 long distance runners and compared them to another 14 control group runners [92]. The foot exercise protocol effectively increased intrinsic foot muscle volume by up to 22% and vertical propulsive forces in these recreational runners. Interestingly, actual strength of the toe flexors did not statistically increase, despite the 36% improvement from pre to post intervention testing. In a longer term follow-up of these runners, the authors also found that the group who had participated in the foot strengthening program had a reduced risk of sustaining a running-related injury [93].

6.3

Areas of future biomechanical research

Our understanding of the structure and function of the foot and ankle muscles and tendons has greatly improved in recent years due to the advances in imaging and data collection technology. However, there is still a lot to be uncovered, particularly around individual contributions of specific structures to normal and abnormal foot function. Most of the morphological

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studies in this chapter are based on static measurements. Dynamic measures of muscle and tendon morphology will better inform our understanding of foot and ankle function. A better understanding of the role that foot and ankle muscles play in injury and lower limb pathologies would help clinicians, footwear designers, coaches, and athletes come up with more informed products and/or practices for the prevention and treatment of certain foot and ankle conditions. Most of our knowledge is based on cross-sectional studies; therefore, we cannot infer many of the cause-and-effect relationships between the muscle and tendon properties and other biomechanical variables or foot pathologies investigated. Some of the studies discussed in this chapter have suggested a cause-and-effect relationship between atrophy of the foot structure and certain pathologies, but it is also possible that atrophy may develop secondary to the conditions, particularly in chronic cases. Whether atrophy develops prior to onset or secondary to extended duration of these injuries requires longitudinal, well designed studies. Furthermore, longitudinal studies may help us gain a better understanding of the age-related decline in strength and muscle atrophy, which could help us identify whether or not this could be prevented or reversed. While the exact causes of the reduction in foot muscle strength with aging is not currently known, numerous factors such as motor unit loss [94], a reduction in physical activity, or history of footwear use may contribute to age-related decline. Longitudinal studies are imperative to understand the origin of foot muscle adaptations. Similarly, inclusion of muscle and tendon morphology in prospective studies investigating risk factors for lower limb injury will help us determine mechanistic links to injury risk. Further research around the specificity of foot strengthening interventions is also warranted to determine whether specific muscles can be retrained. When designing exercises to isolate the intrinsic foot muscles, the aim should be for flexion of the proximal interphalangeal and metatarsophalangeal joints through their respective ranges of motion while limiting motion at the distal interphalangeal joint. However, when trying to address specific pathologies such as a planus foot, it may be worthwhile adding specific exercises that target, for example, TP.

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[22] Granado MJ, et al. Metatarsophalangeal joint extension changes ultrasound measurements for plantar fascia thickness. J Foot Ankle Res 2018;11(1):20. [23] Kelly LA, et al. Intrinsic foot muscles contribute to elastic energy storage and return in the human foot. J Appl Physiol 2019;126(1):231 8. [24] Kelly LA, Lichtwark G, Cresswell AG. Active regulation of longitudinal arch compression and recoil during walking and running. J R Soc Interface 2015;12(102):20141076. [25] Okamura K, et al. The effect of additional activation of the plantar intrinsic foot muscles on foot dynamics during gait. Foot 2018;34:1 5. [26] Kirane YM, Michelson JD, Sharkey NA. Evidence of isometric function of the flexor hallucis longus muscle in normal gait. J Biomech 2008;41 (9):1919 28. [27] Kelly LA, et al. The influence of foot-strike technique on the neuromechanical function of the foot. Med Sci Sports Exerc 2018;50(1):98 108. [28] de Almeida MO, et al. Is the rearfoot pattern the most frequently foot strike pattern among recreational shod distance runners? Phys Ther Sport 2015;16(1):29 33. [29] Davis IS, Rice HM, Wearing SC. Why forefoot striking in minimal shoes might positively change the course of running injuries. J Sport Health Sci 2017;6(2):154 61. [30] Hashizume S, Yanagiya T. Forefoot strike requires higher impulse of the Achilles tendon force than rearfoot strike. Footwear Sci 2015;7(Suppl. 1):S140 1. [31] Dorn TW, Schache AG, Pandy MG. Muscular strategy shift in human running: dependence of running speed on hip and ankle muscle performance. J Exp Biol 2012;215(Pt 11):1944 56. [32] Tanaka T, et al. Characteristics of lower leg and foot muscle thicknesses in sprinters: does greater foot muscles contribute to sprint performance? Eur J Sport Sci 2019;19(4):442 50. [33] Garofolini A, et al. Effect of habitual foot strike patterns on muscle and bone properties of the foot. Footwear Sci 2019;11(Suppl. 1):S142 4. [34] Farris DJ, et al. The functional importance of human foot muscles for bipedal locomotion. Proc Natl Acad Sci 2019;201812820. [35] Fujiwara K, et al. Changes in muscle thickness of gastrocnemius and soleus associated with age and sex. Aging Clin Exp Res 2010;22 (1):24 30. [36] Narici MV, et al. Effect of aging on human muscle architecture. J Appl Physiol 2003;95(6):2229 34. [37] Goodpaster BH, et al. The loss of skeletal muscle strength, mass, and quality in older adults: the health, aging and body composition study. J Gerontol Ser A: Biol Sci Med Sci 2006;61(10):1059 64. [38] Endo M, Ashton-Miller J, Alexander N. Effects of age and gender on toe flexor muscle strength. J Gerontol Ser A, Biol Sci Med Sci 2002;57A (6):M392 7. [39] Menz HB, et al. Plantarflexion strength of the toes: age and gender differences and evaluation of a clinical screening test. Foot Ankle Int 2006;27(12):1103 8. [40] Menz HB, Morris ME, Lord SR. Foot and ankle risk factors for falls in older people: a prospective study. J Gerontol Ser A, Biol Sci Med Sci 2006;61A(8):866 70. [41] Mickle KJ, et al. ISB Clinical Biomechanics Award 2009: toe weakness and deformity increase the risk of falls in older people. Clin Biomech 2009;24:787 91. [42] Wall JR, Harkness MA, Crawford A. Ultrasound diagnosis of plantar fasciitis. Foot Ankle 1993;14(8):465 70. [43] Allen RH, Gross MT. Toe flexors strength and passive extension range of motion of the first metatarsophalangeal joint in individuals with plantar fasciitis. J Orthop Sports Phys Ther 2003;33(8):468 78. [44] Chang R, Kent-Braun JA, Hamill J. Use of MRI for volume estimation of tibialis posterior and plantar intrinsic foot muscles in healthy and chronic plantar fasciitis limbs. Clin Biomech 2012;27(5). [45] Cheung RTH, et al. Intrinsic foot muscle volume in experienced runners with and without chronic plantar fasciitis. J Sci Med Sport 2016;19 (9):713 15. [46] Thordarson DB, et al. Dynamic support of the human longitudinal arch. A biomechanical evaluation. Clin Orthop Relat Res 1995;316:165 72. [47] Wacker J, et al. MR morphometry of posterior tibialis muscle in adult acquired flat foot. Foot Ankle Int 2003;24(4):354 7. [48] Gray EG, Basmajian JV. Electromyography and cinematography of leg and foot (“normal” and flat) during walking. Anat Rec 1968;161(1):1 15. [49] Murley GS, Menz HB, Landorf KB. Foot posture influences the electromyographic activity of selected lower limb muscles during gait. J Foot Ankle Res 2009;2(1):35. [50] Hunt AE, Smith RM. Mechanics and control of the flat vs normal foot during the stance phase of walking. Clin Biomech 2004;19(4):391 7. [51] Zhang X, Aeles J, Vanwanseele B. Comparison of foot muscle morphology and foot kinematics between recreational runners with normal feet and with asymptomatic over-pronated feet. Gait Posture 2017;54:290 4. [52] Angin S, et al. Ultrasound evaluation of foot muscles and plantar fascia in pes planus. Gait Posture 2014;40(1):48 52. [53] Murley GS, et al. Foot posture is associated with morphometry of the peroneus longus muscle, tibialis anterior tendon, and Achilles tendon. Scand J Med Sci Sports 2013; In Print. ¨ , Korkusuz F. Morphological and mechanical properties of plantar fascia and intrinsic foot muscles in individuals with and ¨ nlu¨er NO [54] Ta¸s S, U without flat foot. J Orthop Surg 2018;26(3). p. 2309499018802482. [55] Fukumoto Y, et al. Navicular drop is negatively associated with flexor hallucis brevis thickness in community-dwelling older adults. Gait Posture 2020;78:30 4. [56] Okamura K, et al. The effect of additional activation of the plantar intrinsic foot muscles on foot kinematics in flat-footed subjects. Foot 2019;38:19 23.

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[57] Jung D-Y, Koh E-K, Kwon O-Y. Effect of foot orthoses and short-foot exercise on the cross-sectional area of the abductor hallucis muscle in subjects with pes planus: a randomized controlled trial. J Back Musculoskelet Rehab 2011;24(4):225 31. [58] Dunn JE, et al. Prevalence of foot and ankle conditions in a multiethnic community sample of older adults. Am J Epidemiol 2004;159(5):491 8. [59] Munro BJ, Steele JR. Foot-care awareness. A survey of persons aged 65 years and older. J Am Podiatr Med Assoc 1998;88(5):242 8. [60] van Schie C, et al. Muscle weakness and foot deformities in diabetes: relationship to neuropathy and foot ulceration in Caucasian diabetic men. Diabetes Care 2004;27(7):1668 73. [61] Scott G, Menz HB, Newcombe L. Age-related differences in foot structure and function. Gait Posture 2007;26(1):68 75. [62] Myerson MS, Shereff MJ. Pathological Anat claw hammer toes. J Bone Jt Surg 1989;71-A(1):45 9. [63] Stewart S, et al. Ultrasonic evaluation of the abductor hallucis muscle in hallux valgus: a cross-sectional observational study. BMC Musculoskelet Disord 2013;14(1):1 6. [64] Mickle KJ, Nester CJ. Morphology of the toe flexor muscles in older adults with toe deformities. Arthritis Care Res (Hoboken) 2018;70(6):902 7. [65] Hurn SE, Vicenzino B, Smith MD. Functional impairments characterizing mild, moderate, and severe Hallux Valgus. Arthritis Care Res 2015;67(1):80 8. [66] Arinci Incel, et al. Muscle imbalance in Hallux Valgus: an electromyographic study. Am J Phys Med Rehab 2003;82(5):345 9. [67] Mann RA, Hagy Jl. The function of the toes in walking, jogging and running. Clin Orthop Relat Res 1979;142:24 9. [68] Glasoe W. Treatment of progressive first metatarsophalangeal hallux valgus deformity: a biomechanically based muscle-strengthening approach. J Orthop Sports Phys Ther 2016;46(7):596 605. [69] Kirkman MS, et al. Diabetes in older adults. Diabetes Care 2012;35(12):2650 64. [70] Andersen H, Gjerstad MD, Jakobsen J. Atrophy of foot muscles: a measure of diabetic neuropathy. Diabetes Care 2004;27(10):2382 5. [71] Cheuy VA, et al. Intrinsic foot muscle deterioration is associated with metatarsophalangeal joint angle in people with diabetes and neuropathy. Clin Biomech 2013;28(9 10):1055 60. [72] Andersen H, et al. Muscular atrophy in diabetic neuropathy: a stereological magnetic resonance imaging study. Diabetologia 1997;40(9):1062 9. [73] Andreassen CS, et al. Accelerated atrophy of lower leg and foot muscles—a follow-up study of long-term diabetic polyneuropathy using magnetic resonance imaging (MRI). Diabetologia 2009;52(6):1182 91. [74] Severinsen K, et al. Atrophy of foot muscles in diabetic patients can be detected with ultrasonography. Diabetes Care 2007;30(12):3053 7. [75] Bus SA, et al. Intrinsic muscle atrophy and toe deformity in the diabetic neuropathic foot: a magnetic resonance imaging study. Diabetes Care 2002;25(8):1444 50. [76] Holowka NB, Wallace IJ, Lieberman DE. Foot strength and stiffness are related to footwear use in a comparison of minimally—vs. conventionally-shod populations. Sci Rep 2018;8(1). p. 3679-3679. [77] Potthast W, et al. The choice of training footwear has an effect on changes in morphology and function of foot and shank muscles. Beijing, China: International Society of Biomechanics in Sport (ISBS); 2005. [78] Potthast W, et al. Changes in morphology and function of toe flexor muscles are related to training footwear. In: 7th Symposium on footwear biomechanics. Cleveland, OH: Case Western Reserve University Printing Services; 2005. [79] Bruggemann, G.-P., et al. Effect of increased mechanical stimuli on foot muscles functional capacity. In: ISB XXth congress. Cleveland, OH; 2005. [80] Zhang X, et al. The morphology of foot soft tissues is associated with running shoe type in healthy recreational runners. J Sci Med Sport 2018;21(7):686 90. [81] Ridge ST, et al. Foot bone marrow edema after a 10-wk transition to minimalist running shoes. Med Sci Sports Exerc 2013;45(7):1363 8. [82] Miller EE, et al. The effect of minimal shoes on arch structure and intrinsic foot muscle strength. J Sport Health Sci 2014;3(2):74 85. [83] Csapo R, et al. On muscle, tendon and high heels. J Exp Biol 2010;213(Pt 15):2582 8. [84] Akuzawa H, et al. Calf muscle activity alteration with foot orthoses insertion during walking measured by fine-wire electromyography. J Phys Ther Sci 2016;28(12):3458 62. [85] Murley GS, Landorf KB, Menz HB. Do foot orthoses change lower limb muscle activity in flat-arched feet toward a pattern observed in normal-arched feet? Clin Biomech 2010;25(7):728 36. [86] Kokkonen J, et al. Improved performance though digit strength gains. Res Q Exerc Sport 1988;59(1):57 63. [87] Unger CL, Wooden MJ. Effect of foot intrinsic muscle strength training on jump performance. J Strength Cond Res 2000;14(4):373 8. [88] Goldmann J-P, et al. Effects of increased toe flexor muscle strength to foot and ankle function in walking, running and jumping. Footwear Sci 2011;3(S1):S59 60. [89] Kobayashi R, et al. Effects of toe grasp training for the aged on spontaneous postural sway. J Phys Ther Sci 1999;11(1):31 4. [90] Kim MH, et al. Comparison of muscle activities of abductor hallucis and adductor hallucis between the short foot and toe-spread-out exercises in subjects with mild hallux valgus. J Back Musculoskelet Rehabil 2013;26(2):163 8. [91] Kim M-H, et al. Effect of toe-spread-out exercise on hallux valgus angle and cross-sectional area of abductor hallucis muscle in subjects with hallux valgus. J Phys Ther Sci 2015;27(4):1019 22. [92] Taddei UT, et al. Effects of a foot strengthening program on foot muscle morphology and running mechanics: a proof-of-concept, single-blind randomized controlled trial. Phys Ther Sport 2020;42:107 15. [93] Taddei UT, et al. Foot core training to prevent running-related injuries: a survival analysis of a single-blind, randomized controlled trial. Am J Sports Med 2020;48(14):3610 19. [94] Doherty TJ, et al. Effects of motor unit losses on strength in older men and women. J Appl Physiol 1993;74(2):868 74.

Chapter 7

Ligaments Aerie Grantham1,2, Joseph M. Iaquinto1,2,3 and Alexander Berardo-Cates1 1

RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 2Department of

Mechanical Engineering, University of Washington, Seattle, WA, United States, 3Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Abstract As with all structural tissues of the body, ligaments play an important functional role in biomechanics. However, the small size and interwoven complexity of ligaments of the foot and ankle has traditionally made research on their behavior much more difficult to study than other, larger ligaments of the body, such as the anterior cruciate ligament. This chapter details several key ligaments and ligament groups, noting their anatomy and function, and the key literature that describes their mechanical and physiological behavior. The effects of age, disease, and common pathologies on ligament tissue are briefly reviewed as to their impact on tissue biomechanics. Additionally, the unique morphology of each ligament as well as their viscoelastic nature, tissue environment conditions, and complex geometry present challenges to the evaluation of mechanical properties via traditional mechanical testing fixation strategies. This chapter provides an overview of prior work done to determine foot and ankle ligament mechanical properties, with special emphasis on challenges facing these investigations, and potential directions toward overcoming them.

7.1

Introduction

Every structural tissue in the human body plays important functional roles that support our survival and movement. Muscles move bones via tendons, bones serve as the anchoring foundation for our tissues and organs, cartilage cushions the motion between bones, and ligaments play important roles in stabilizing and constraining the motions of our joints. The complexity of the local ligament network is dependent somewhat on the number and proximity of such joints. The foot and ankle is comprised of many unique bones and bony articulations with an equally varied set of ligaments, each contributing to joint mobility and stability. Ligaments are viscoelastic, like almost all tissues of the human body; and studies to capture this behavioral complexity require specialized tests and considerations. Investigating the mechanical properties of ligaments is a fundamental task that can help us understand the form and function of the foot and ankle. Knowledge of their behavior in health and injury better equips clinicians to help their patients. Ankle ligament injuries are commonly seen by primary care providers and emergency departments and account for 50% of all sports-related injuries [1]; this has cause for concern as ankle injuries and sprains may lead to chronic instability and pain [2]. The ligaments of the foot help to maintain the medial, lateral, and transverse arches of the foot to support the weight of the body. When repeated stretching or traumatic injury damages these ligaments, severe pain can result. In some cases, this can even contribute to limited mobility due to chronic injury. One common example of this is plantar fasciitis, where the plantar fascia is injured, with an incidence of more than one million cases per year in the United States [3]. This warrants careful study of the foot and ankle ligaments and their mechanical properties.

7.2

Ligament anatomy

The foot and ankle is comprised of 28 bones and numerous articulations between them. Most bone-to-bone articulation is supported by a joint capsule with thickened reinforcements at key locations called ligaments. Depending on which joint they cross, their length, and the motions they limit, some ligaments are thick and cord-like while others are thin Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00046-9 © 2023 Elsevier Inc. All rights reserved.

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and more membrane-like. Because of the sheer number of ligaments present in the foot and ankle (over 100), their small size and even smaller and intertwined insertions sites, isolating, defining, and testing them all presents a challenge to researchers. This chapter focuses on a few specific regions and structures, but will span the ankle joint, hindfoot, midfoot, and forefoot. For the ankle joint, attention will be given to both the lateral and medial ligament complexes. In the hindfoot, the ligaments stabilizing the subtalar joint will be discussed. The midfoot review will primarily focus on the plantar fascia but consideration will be given to the dorsal, interosseous, and other plantar ligaments present that support the base of the metatarsals. Finally, the forefoot section will briefly discuss the deep transverse metatarsal ligaments than span across all the distal heads of the metatarsals and metatarsophalangeal joints.

7.3

Mechanical properties

Each ligament of the foot and ankle has a characteristic size, shape, insertion, and loading requirement, which deserves special consideration when determining functional mechanical properties. Ligaments gradually calcify near their bony insertions, altering their mechanical response. This impacts in vivo ligament failure as the ligament can fail (tear) between insertion sites, at the ligament insertion, or the ligament and its bony anchor can separate from the main body of the bone (i.e., avulsion). Since the variation in ligament anatomy doesn’t conform to an idealized engineered coupon, applying standardized mechanical tests to determine ligament mechanical property data can see high variability between studies. Over the years, researchers have examined different combinations of ligaments in the foot and ankle, as well as specific ligaments in isolation. A representative selection of studies and their results regarding ligament mechanical properties for some specific structures of the foot and ankle is provided in the following sections.

7.3.1 Ankle joint 7.3.1.1 Lateral collateral ligaments Excessive foot inversion is limited in part by the lateral collateral ligament (LCL) complex (Fig. 7.1, Tables 7.1 and 7.2), which radiates from the lateral malleolus and inserts onto the nearby tibia, talus, and calcaneus. The LCL consists of the anterior talofibular ligament (ATFL), a thin flat ligament with disorganized fibers, which inserts between the distal fibula across to the lateral talar body. The ATFL has the lowest reported ultimate failure load in this complex. The calcaneofibular ligament (CFL) is a cord-like ligament that has highly organized fibers. This ligament inserts near the AFTL on the distal fibula and crosses a greater distance to the body of the posterior lateral calcaneus. The posterior talofibular ligament (PTFL) inserts opposite of the ATFL on the distal fibula and runs very posterior to the talus. The CFL and PTFL have similar strength characteristics. The lower ultimate load of the ATFL, along with its fiber orientation and anterior position, may explain why this ligament sees frequent injury [4].

FIGURE 7.1 LCL complex [5]. Lateral view of the ankle joint, mid-foot forward omitted (right side of image), with bones and ligament structures labeled. LCL, lateral collateral ligament.

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TABLE 7.1 Comparison of mean LCL stiffnesses identified from literature. Stiffness (kN/m) Study

ATFL

CFL

PFTL

Attarian [6]

39.99 6 8.54

7.05 6 0.69

3.98 6 1.38

Siegler [4]

141.8 6 79.3

126.6 6 42.9

164.3 6 55.5

Rochelle [5]

44.7 6 16.6

45.8 6 19

59.0 6 10.7

LCL, lateral collateral ligament

TABLE 7.2 Comparison of mean LCL ultimate loads identified from literature. Ultimate load (N) Study

ATFL

CFL

PFTL

Attarian [6]

138.9 6 23.5

345.7 6 55.2

261.2 6 32.4

Siegler [4]

231 6 129

307 6 142

418 6 191

Rochelle [5]

263.6 6 164.3

367.8 6 79.8

351.4 6 10.7

LCL, lateral collateral ligament.

FIGURE 7.2 Ligaments on the medial aspect of the ankle, calcaneus to the left and midfoot to the right. (A) tibionavicular, (B) anterior tibiotalar, (C) tibiospring portion of the tibiocalcaneal, (D) tibiocalcaneal, and (E) posterior tibiotalar [7].

7.3.1.2 Medial collateral ligaments The medial collateral ligament (MCL) complex (Fig. 7.2), or the deltoid ligament, assists in limiting ankle eversion. This complex includes the tibionavicular ligament, a broad and fan-like structure coursing from the anterolateral aspect of the tibia to the navicular. This structure has the lowest measured strength of the MCL complex with an ultimate failure load of 120.0 N [4]. Also included is the posterior tibiotalar ligament, a thick ligament with highly organized fibers that courses from the distal tibia to the medial and posterior talus, and has high elastic stiffness, high yield load, and high ultimate load. These properties in conjunction with its fiber orientation and posterior location contributes to stabilizing the talus during ankle eversion and ankle dorsiflexion [4]. Other components include the anterior tibiotalar and the tibiocalcaneal (a portion known as the tibiospring).

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FIGURE 7.3 Schematic of the subtalar ligaments. (A) Sagittal plane with the anterior capsular ligament (ACL), interosseous talocalcaneal ligament (ITCL), and cervical ligament (CL). (B) Coronal plane with the same ligaments [9]. Image used from work distributed under the terms of the Creative Commons Attribution 4.0 International license http://creativecommons.org/licenses/by/4.0/.

7.3.2 Hindfoot ligaments 7.3.2.1 Subtalar joint The talocalcaneal or subtalar joint accommodates coronal plane movement, which helps the foot conform to uneven ground. Within the subtalar joint region, the anterior capsular ligament and interosseous talocalcaneal ligament (ACLITCL) complex is stiffer (150 6 51 kN/m) than the cervical ligament (CL) and has a higher reported failure load (382.5 6 158.0 N) [8]. Although short and thin, the ITCL extends the width of the ankle deep between the bones where it inserts around articular facets between the talus and calcaneus. These strength properties are ideal for the ACL-ITCL complex’s function of maintaining apposition of the subtalar joint while the lower stiffness of the CL may aid the subtalar joint in adapting to the ground throughout gait [8] (Fig. 7.3).

7.3.2.2 Talonavicular joint The spring ligament’s main role is to stabilize the talonavicular joint and support the medial longitudinal arch. This complex supports the head of the talus as well as connecting the calcaneus and navicular. This support is due to its anatomy, which is comprised of three bundles referred to as the superomedial, medioplantar oblique, and inferomedial calcaneonavicular ligament [10]. However much like the midfoot ligaments in the following section, the complex morphology of this structure makes it difficult isolate and clearly distinguish the bundles; as such, the medioplantar bundle is not always identified as separate from the others. Determining the mechanical properties for the spring ligament is crucial to better understanding injury mechanisms of this complex, namely what its role is in the development of flatfoot deformity [11].

7.3.3 Midfoot ligaments 7.3.3.1 Plantar fascia The plantar aponeurosis or fascia (Fig. 7.4), a broad and thick connective tissue which supports the medial arch of the foot, has demonstrated little strain rate dependence [12,13]. When the medial, lateral, and middle portions of the plantar fascia are tested as one structure, an average stiffness of 203.7 6 50.5 kN/m has been found for all loading rates. A recent work performed by Isvilanonda [12] has evaluated the mechanical properties of each region individually by selectively resecting regions and measuring the mechanical properties of each separately (Fig. 7.5). The mechanical properties differ by region (Table 7.3), as the medial and lateral regions had a significantly lower modulus when compared against the proximal and middle distal regions of the tissue [12].

7.3.3.2 Metatarsal base ligaments The midfoot region contains numerous short ligaments that give rise to a complex network making not only testing difficult but also clearly defining ligaments a challenge. Discussed here are the results from two studies examining the mechanical properties of midfoot ligaments (Table 7.4). In the dorsal region, the dorsal cuneometatarsal ligament (connecting the first cuneiform and second metatarsal) is significantly weaker than the interosseous ligament (often referred to as the Lisfranc ligament) [14 16] and the plantar tarsometatarsal ligament (specifically the ligament connecting the first cuneiform to the second and third metatarsal) [14]. The interosseous ligament has, in some studies, been grouped

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FIGURE 7.4 Bone-plantar fascia-bone lateral view from a right foot with calcaneus imbedded in a resin block [12].

FIGURE 7.5 Specimen with medial and lateral bands resected [12].

together with plantar ligaments as part of a bundle instead of being differentiated. When comparing the interosseous ligament to the plantar tarsometatarasal ligament, the interosseous is both significantly stronger and stiffer, suggesting that any injuries to this ligament may be more destabilizing than an injury to one of the other midfoot ligaments [14].

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TABLE 7.3 Average stiffness for plantar fascia region [12]. Region

Stiffness (N/mm)

Elastic modulus (MPa)

Proximal middle

442 6 104.3

400 6 59

Medial

576 6 89.4

225 6 66

Distal middle

1207 6 328.6

522 6 62

Lateral

166 6 35.0

242 6 24

TABLE 7.4 Midfoot ligament mechanical properties [14,15]. Study

Ligament

Ultimate load

Stiffness (kN/m)

Solan et al. [14]

Dorsal cuneometatarsal

170 6 33

40 6 9

Interosseous

449 6 58

90 6 3

Plantar tarsometatarsal

305 6 38

62 6 3

Dorsal cuneometatarsal

150.7 6 33.1

Interosseous

368.8 6 126.8

Kura et al. [15]

Ultimate Stress (MPa)

Ultimate strain (%)

Modulus of elasticity (MPa)

66.3 6 18.3

5.9 6 2.3

115.6 6 34.9

5.5 6 2.7

189.7 6 57.2

5.8 6 29.2

87.0 6 29.2

7.4 6 3.6

7.3.4 Forefoot Spanning across all metatarsals is the deep transverse metatarsal ligament, which provides intermetatarsal support. Abdalbary et al. tested a single transverse metatarsal band situated between the first and second metatarsal [17]. This ligament, which experienced a maximum stress of 13.3 6 6 MPa, stabilizes the first MTP joint by resisting dorsiflexion and dorsal dislocations. Further investigation of the mechanical properties of this ligament may provide insight to potential corrective actions needed to treat hallux valgus [17].

7.3.5 Variations in mechanical properties 7.3.5.1 Changes in activity level The structural and mechanical properties of ligaments vary between individuals and populations based on their response to loading conditions, and the presence or absence of disease. Like many other tissues of the body, ligament properties change in response to activity level. Previous work has shown via testing with animal ligament tissues that endurance type exercises can increase the stiffness and strength of the tissue [18]. However, a lack of stress on ligaments over a period of 6 weeks or more will yield a decrease in mechanical strength [19]. Therefore, people who suffer from injuries or conditions that restrict range of motion or prohibit foot and ankle loading will have ligaments that are weaker when compared to a healthy population.

7.3.5.2 Foot comorbidity Foot comorbidities pose a potential risk to the integrity of ligaments and their mechanical properties. When compared to healthy feet, ligaments from feet afflicted with hallux valgus saw significantly lower maximum force, stress, and strain [17]. Diabetes also warrants concern for ligament properties. Collagen is negatively affected by diabetes and because collagen fibers are significant contributors to ligament strength properties, optimal stiffness and strength can no longer be achieved and this poses a greater risk for injury [20].

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7.3.5.3 Age effects Aging populations are at a greater risk of ligament injuries and joint instability than younger populations. Ligament mechanical properties decrease significantly between younger and elderly populations. There are numerous factors that lead to this degradation, including decreased physical activity and loading level of the ligaments as well as changes in joint geometry with age. In the elderly population, ligaments have been found to fail more often at the mid substance rather than the insertion sites, suggesting it is the mid substance of the ligament that degrades more rapidly [21].

7.3.5.4 Influence of anthropometric effects When considering anthropometric effects, previous work has shown that ligaments from males generally have a larger cross-sectional area as well as significantly higher yield and ultimate strength when compared to females [4]. Ligaments from donors with higher body mass index (BMI) have higher ultimate failure loads than donors with a lower BMI [5]. From these findings we can conclude that the male populations have larger and stronger ligaments due to their generally larger size/BMI, and therefore increased mechanical loading, as compared to the female population.

7.4

Ligament sprains

Injuries to ligaments not only affect their form but also their function, providing stability to joints. As previously mentioned, severe ankle sprains and/or numerous repeated sprains can lead to chronic ankle instability. The greatest risk of an ankle sprain is history of a previous ankle sprain. The LCL complex is typically injured during a sprain. Excessive inversion of the foot with external leg rotation is the most common mechanism of a sprain [1]. However, in high energy activities or collision sports, high ankle sprain of the anterior and posterior inferior tibiofibular ligament is common. Treatment of sprains often includes rest, icing the affected area, compression, and elevating the foot. Following initial treatment is physical rehabilitation (e.g., strength training) before progressing to normal unrestrained activity. In the case of conservative treatment management being insufficient or leading to recurrent ankle instability due to ligament injury, various surgical procedures might be recommended [1]. Ligament instability due to severe and chronic sprain generally falls into two broad categories: anatomic repair and nonanatomic repair. The former attempts to reconstruct the normal anatomy by shortening and suturing ligament midsubstance or other soft tissue such as a neighboring retinaculum. Nonanatomic reconstructions use tissue grafts (of primarily tendon) that are tunneled through and sutured to bone. Outcomes from surgical interventions are based on several factors, some of which include patient activity, anatomical alignment, and tissue physiology [22].

7.5

Overcoming limitations

The composition and size of ligament tissue exhibit several challenges to traditional fixation and mechanical testing methodologies. Approximately 80% of a ligament’s dry weight is type I collage fibers, but often the arrangement of those fibers is not unidirectional to better accommodate loading from multiple axes [23]. Because of this variability in fiber direction, it is challenging to align fibers with applied testing loads, which challenges testing repeatability and data aggregation. One approach is to maintain anatomical position during testing to better replicate in vivo conditions. Another approach is to cut dog-bone shaped specimens to adapt established testing protocols developed by the American Society for Testing and Materials (ASTM) [24]. Because the tissue is being cut from the bony insertions and clamped to the testing fixture directly, the properties of the ligament insertion are not considered. However, the use of ASTM testing standards might improve testing repeatability and the ability to generalize mechanical properties. Results from the three aforementioned studies evaluating the LCL properties (Tables 7.1 and Table 7.2) highlight the variability that is common in this area of research (this is especially evident in Table 7.1). There are several factors that contribute to this, one of which is the differences in methodology between these studies. For example, Siegler [4] encased the bony attachments in resin (potting) before mounting to the test fixture while the other two works attached the bony attachments to their fixture without potting. Due to the stress-relaxation response of viscoelastic materials, the testing methodologies for ligaments must be repeatable across multiple strain rates. Review work evaluating the effects of strain rate on ligament mechanical properties showed that strain rate has minimal effect on ligament mechanical properties. Therefore, it is common to precondition tissues through cyclic loading to stabilize hysteresis so that the time dependent responses are repeatable during nonlinear elastic measurements [25].

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To acquire comprehensive mechanical property data, the testing protocol will require time and effort being spent to best preserve the integrity of the ligament tissue being tested. While working with specimens, it is critical to keep the tissues hydrated throughout testing as previous studies have shown that properties are altered merely by the interaction between the isotonic fluid media and ligament tissue [26]. Methods for maintaining hydration are varied such as wrapping ligaments in Ringer’s solution and frequent hydration via 10% saline solution [4,6]. A recent work has maintained hydration automatically by spraying the ligament tissue with saline solution at specific intervals within a specialized environment chamber during testing (Fig. 7.6) [27]. Another study implemented an environmental chamber that could simulate physiological conditions like temperature and near 100% humidity (Fig. 7.7) [12]. Both methods require minimal intervention on the user’s part to improve consistency and repeatability of the time-consuming mechanical tests.

FIGURE 7.6 Automated testing setup equipped with a controlled environmental chamber [27].

FIGURE 7.7 Environmental chamber custom designed to keep the interior near physiological conditions [12].

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TABLE 7.5 Structural properties before and after refreezing the femur-medial collateral ligament-tibia complex [28]. Structural properties

Fresh

Refrozen

Stiffness (N/mm)

106.5 6 5.9

107.7 6 11.2

Ultimate load (N)

331.0 6 81.8

336.7 6 50.1

Elongation at Failure (mm)

4.2 6 0.9

4.3 6 0.8

Energy absorbed (N mm)

703.8 6 270.1

697.2 6 264.9

TABLE 7.6 Mechanical properties before and after refreezing the medial collateral ligaments [28]. Mechanical properties

Fresh

Refrozen

Tangent Modulus (MPa)

1107.2 6 126.3

1056.2 6 207.9

Ultimate tensile strength (MPa)

84.4 6 22.2

83.2 6 16.0

Ultimate strain (%)

10.6 6 2.8

11.1 6 4.1

Strain energy density (MPa)

4.62 6 2.35

4.58 6 2.53

FIGURE 7.8 Comparison between fresh and stored specimens; freezing ligaments does increase the peak stress during cyclic loading [29].

Additionally, freezing specimens to preserve their integrity in between dissection and testing is often required. Refreezing and thawing have the potential to affect mechanical properties, though previous work has indicated that the effects on peak stress are not statistically significant as long as caution and tissue hydration are considered especially for repeated freezing and thawing cycles (Tables 7.5 and 7.6) [28]. Freezing ligaments does increase the peak stress during cyclic loading (Fig. 7.8) and reduces the hysteresis area for the initial ten cycles by a statistically significant amount (Fig. 7.9) [29]. It is challenging to secure ligaments to testing fixtures without damaging them due to their short length and complex morphology. Previous work has potted the ligament’s bony attachments in plastic resin. The resin block is then secured to the testing fixture (Fig. 7.10) [4,30]. Ligaments have also been attached by leaving the insertion areas intact and mounting the bone to the test fixture without potting in resin [6]; testing has also been performed by dissecting the bony attachments and affixing them within a custom fixture [5]. As methodology differs across works, it warrants consideration when comparing results and conclusions between studies.

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FIGURE 7.9 Comparison of hysteresis between fresh and stored samples shows significant reduction in the hysteresis area for both the initial and tenth cycles [29].

FIGURE 7.10 Ligament bony attachments potted in a plastic resin and clamped in supports [30].

To measure a true stress response the cross-sectional area must be measured but the irregular morphology of these ligaments introduces challenges when making this measurement. Calipers and micrometers are the tools commonly used across various studies (Fig. 7.11), but lack the accuracy of noncontact methods such as using a microscope, laser telemetric system, or the optical method [31]. However, higher accuracy methods often require highly specialized and expensive equipment. Previous work has developed a method of obtaining ligaments’ geometric features using a commercially available 3D scanner at a resolution that is in agreement with the aforementioned noncontact techniques (Fig. 7.12) [31]. Molding techniques have been used with some success, but this technique can come with a myriad of challenges such as: material shrinkage, failure to capture smaller details and concavities especially with the unique and varied shape of ankle and foot ligaments, difficulty determining the cross-section perpendicular to the body surface, and the need to destroy the mold to get the ligament out of the material [32]. A molding and casting method was developed with materials that minimize casting material shrinkage. This approach used liquid polyurethane and reusable molds made of silicone to accurately capture ligament surface geometry [32]. In recent years, a novel method has been

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FIGURE 7.11 Length measurement of the middle fibers of the deltoid ligament [27].

FIGURE 7.12 3D scan system with turntable, camera, and calibration pattern used to determine ACL geometry [31].

developed to determine ligament length and cross-sectional areas using segmented weight-bearing foot CT images (Fig. 7.13) and an iterative refinement process to obtain accurate measurements [27]. Ligaments that wrap around bony features before reaching their terminal insertion provide unique challenges in replicating loading conditions and mechanical response. To overcome these challenges, a recent study has mounted intact cadaveric legs into a custom mechanical testing device and measured load and displacement responses at the time of ligament rupture [33]. The timing of ligament injury was determined from acoustic sensors, strain gauges, force and moment measurements, and 3D bone kinematics. Innovations in elastography have provided opportunities to measure the elastic and stiffness properties of soft tissues in vivo [34,35]. These techniques have been applied to large superficial ligaments such as those in the knee [36,37], and ankle [38,39] with some success. However, these methods face limitations in their ability to validate results against a ground truth due to the ethical restraint of testing in vivo human ligament specimens, as no noninvasive validation method is currently available. Ligament strain elastography validation using cadaveric in situ specimens under physiologic loading conditions is currently an active research area [37].

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FIGURE 7.13 Ligament 3D geometry over bony anatomy [27].

7.6

Future areas of research

Nonpathological ligaments in the foot and ankle have mechanical properties necessary to maintain the stability of the joints of the foot and ankle. However, the viscoelastic nature and complex morphology of these ligaments require comprehensive testing and measurement capturing methodologies that are often both expensive and time intensive. Determining ankle and foot ligament mechanical properties is crucial in understanding the biomechanical function, the injuries that can happen, and the best path toward injury prevention or treatment. The limitations and methods researchers have devised to overcome them highlight the difficulty that comes with testing foot and ankle ligaments. Some methodologies, such as ultrasound elastography, show promise in measuring in vivo ligament function, though work toward validated standards remains. It is clear that a standard procedure for making these measurements does not exist, and there is plenty of room to develop novel innovations. Fully characterizing ligament behavior is essential when working toward validated computational simulations of joint biomechanics [25].

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[14] Solan MC, Moorman CT, Miyamoto RG, Jasper LE, Belkoff SM. Ligamentous restraints of the second tarsometatarsal joint: a biomechanical evaluation. Foot Ankle Int. 2001;22(8):637 41. Available from: https://doi.org/10.1177/107110070102200804. [15] Kura H, Luo Z-P, Kitaoka HB, Smutz WP, An K-N. Mechanical behavior of the lisfranc and dorsal cuneometatarsal ligaments: in vitro biomechanical study. J Orthop Trauma 2001;15(2):107 10 February. [16] Sripanich Y, Steadman J, Kra¨henbu¨hl N, Rungprai C, Saltzman CL, Lenz AL, et al. Anatomy and biomechanics of the lisfranc ligamentous complex: a systematic literature review. J Biomech 2021;119:110287. Available from: https://doi.org/10.1016/j.jbiomech.2021.110287 ISSN 0021-9290. [17] Abdalbary SA, Elshaarawy E, Khalid B. Tensile properties of the deep transverse metatarsal ligament in hallux valgus: a CONSORT-compliant article. Medicine 2016;95(8):e2843. Available from: https://doi.org/10.1097/MD.0000000000002843. [18] Cabaud HE, Chatty A, Gildengorin V, Feltman RJ. Exercise effects on the strength of the rat anterior cruciate ligament. Am J Sports Med 1980;8(2):79 86. Available from: https://doi.org/10.1177/036354658000800204. [19] Keira M, Yasuda K, Kaneda K, Yamamoto N, Hayashi K. Mechanical properties of the anterior cruciate ligament chronically relaxed by elevation of the tibial insertion. J Orthop Res 1996;14(1):157 66. Available from: https://doi.org/10.1002/jor.1100140125. Available from: 8618159. [20] Stolarczyk A, Sarzy´nska S, Gondek A, Cudnoch-Je˛drzejewska A. Influence of diabetes on tissue healing in orthopaedic injuries Clin Exp Pharmacol Physiol 2018;45(7):619 27. Available from: https://doi.org/10.1111/1440-1681.12939Epub 2018 Apr 25. Available from: 29570835. [21] Woo SL, Hollis JM, Adams DJ, Lyon RM, Takai S. Tensile properties of the human femur-anterior cruciate ligament-tibia complex. The effects of specimen age and orientation. Am J Sports Med 1991;19(3):217 25. Available from: https://doi.org/10.1177/036354659101900303. Available from: 1867330. [22] Baumhauer JF, O’Brien T. Surgical considerations in the treatment of ankle instability. J Athl Train 2002;37(4):458 62 PMID: 12937567; PMCID: PMC164377. [23] Hudson DM, Archer M, Rai J, Weis MA, Fernandes RJ, Eyre DR. Age-related type I collagen modifications reveal tissue-defining differences between ligament and tendon. Matrix Biol Plus 2021;12:100070. Available from: https://doi.org/10.1016/j.mbplus.2021.100070 ISSN 2590-0285. [24] Schmidt EC, Chin M, Aoyama JT, Ganley TJ, Shea KG, Hast MW. Mechanical and microstructural properties of pediatric anterior cruciate ligaments and autograft tendons used for reconstruction. Orthop J Sports Med 2019;. Available from: https://doi.org/10.1177/ 2325967118821667. [25] Weiss J, Gardiner J. Computational modeling of ligament mechanics. Crit Rev Biomed Eng 2001;29:303 71. Available from: https://doi.org/ 10.1615/CritRevBiomedEng.v29.i3.20. [26] Chimich D, Shrive N, Frank C, Marchuk L, Bray R. Water content alters viscoelastic behaviour of the normal adolescent rabbit medial collateral ligament. J Biomech 1992;25(8):831 7. Available from: https://doi.org/10.1016/0021-9290(92)90223-N ISSN 0021-9290. [27] Berardo-Cates AT. Development of methods for characterizing ankle ligament viscoelastic properties [Master’s thesis]. University of Washington. ResearchWorks; 2019. ,https://digital.lib.washington.edu/researchworks/handle/1773/45230.. [28] Moon DK, Woo SL-Y, Takakura Y, Gabriel MT, Abramowitch SD. The effects of refreezing on the viscoelastic and tensile properties of ligaments. J Biomech 2006;39(6):1153 7. Available from: https://doi.org/10.1016/j.jbiomech.2005.02.012 ISSN 0021-9290. [29] Woo SL, Orlando CA, Camp JF, Akeson WH. Effects of postmortem storage by freezing on ligament tensile behavior. J Biomech 1986;19 (5):399 404. Available from: https://doi.org/10.1016/0021-9290(86)90016-3. Available from: 3733765. [30] Pioletti DP, Rakotomanana LR, Leyvraz P-F. Strain rate effect on the mechanical behavior of the anterior cruciate ligament bone complex. Med Eng Phys 1999;21(2):95 100. Available from: https://doi.org/10.1016/S1350-4533(99)00028-4 ISSN 1350-4533. [31] Hashemi J, Chandrashekar N, Cowden C, Slauterbeck J. An alternative method of anthropometry of anterior cruciate ligament through 3-D digital image reconstruction. J Biomech 2005;38(3):551 5. Available from: https://doi.org/10.1016/j.jbiomech.2004.04.010 ISSN 0021-9290. [32] Schmidt KH, Ledoux WR. Quantifying ligament cross-sectional area via molding and casting. ASME. J Biomech Eng. 2010;132(9):091012. Available from: https://doi.org/10.1115/1.4001881 2010. [33] Mait AR, Forman JL, Nie B, Donlon JP, Mane A, Forghani AR, et al. Propagation of syndesmotic injuries during forced external rotation in flexed cadaveric ankles Orthop J Sports Med. 2018;6(6). Available from: https://doi.org/10.1177/23259671187813332325967118781333. Available from: 30090832. [34] Wells PN, Liang HD. Medical ultrasound: imaging of soft tissue strain and elasticity. J R Soc Interface 2011;8(64):1521 49. Available from: https://doi.org/10.1098/rsif.2011.0054 Epub 2011 Jun 16. PMID: 21680780. [35] Sarvazyan A, Hall TJ, Urban MW, Fatemi M, Aglyamov SR, Garra BS. An overview of elastography—an emerging branch of medical imaging. Curr Med Imaging Rev. 2011;7(4):255 82. Available from: https://doi.org/10.2174/157340511798038684 PMID: 22308105.

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[36] Slane LC, Slane JA, D’hooge J, Scheys L. The challenges of measuring in vivo knee collateral ligament strains using ultrasound. J Biomech 2017;61:258 62. Available from: https://doi.org/10.1016/j.jbiomech.2017.07.020 Epub 2017 Jul 31. PMID: 28802742. [37] Gijsbertse K, Sprengers A, Naghibi Beidokhti H, Nillesen M, de Korte C, Verdonschot N. Strain imaging of the lateral collateral ligament using high frequency and conventional ultrasound imaging: an ex-vivo comparison. J Biomech 2018;73:233 7. Available from: https://doi.org/ 10.1016/j.jbiomech.2018.03.035 Epub 2018 Mar 29. PMID: 29628130. [38] Rougereau G, Marty-Diloy T, Vigan M, Donadieu K, Hardy A, Vialle R, et al. A preliminary study to assess the relevance of shear-wave elastography in characterizing biomechanical changes in the deltoid ligament complex in relation to ankle position Foot Ankle Int 2022;. Available from: https://doi.org/10.1177/1071100722107982910711007221079829. Available from: 35373593. [39] Takaba K, Takenaga T, Tsuchiya A, Takeuchi S, Fukuyoshi M, Nakagawa H, et al. Plantar flexion with inversion shows highest elastic modulus of calcaneofibular ligament using ultrasound share wave elastography J Ultrasound. 2022;. Available from: https://doi.org/10.1007/s40477-02200687-yEpub ahead of print. Available from: 35513766.

Chapter 8

Plantar Soft Tissue William R. Ledoux1,2,3 1

Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States, 2RR&D Center for Limb Loss and MoBility

(CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 3Department of Mechanical Engineering, University of Washington, Seattle, WA, United States

Abstract The plantar soft tissue is a unique structural tissue consisting of adipose and elastic septae, which is surrounded by skin; it serves as the primary mechanical interface between the body and the ground. In this chapter, the gross anatomy is reviewed before the microscopic/histological/biochemical characteristics are considered, and various medical imaging studies are reviewed. The primary functions of the plantar soft tissue are discussed, namely, shock reduction, energy absorption, and load distribution. Several aspects of the mechanical properties of the plantar soft tissue are considered, including the differences between structural in vivo and structural ex vivo testing, as well as material ex vivo results. The findings from ultrasound tests and several other in vivo testing techniques (inverse finite element modeling, magnetic resonance imaging, hand-held force gages, and single plane fluoroscopy) are discussed. Next, research into two important considerations that can affect the mechanical characteristics of the plantar soft tissue (aging and diabetes) are reviewed. Finally, areas of future research are considered.

8.1

Introduction

Between the bones of the foot and the plantar surface are numerous soft tissue layers, including muscle, fat, and skin (epidermis and dermis), all of which is technically “plantar soft tissue.” However, in this chapter, plantar soft tissue refers to the adipose tissue that resides either between skin and bone or skin and muscle. The plantar soft tissue consists of adipocytes enclosed within elastic septal walls, not unlike a closed cell foam. While the cells themselves can deform, the contents of the cells (fat) does not move between cells. This tissue is found throughout the plantar surface of the foot, inferior to the calcaneus, the foot arch, the lateral foot, the metatarsal heads, and even the toes.

8.2

Anatomy

The anatomy of the plantar soft tissue is considered from three aspects: (1) macroscopic or gross anatomy, (2) microscopic or histological or biochemical, and (3) medical imaging, primarily of tissue thickness.

8.2.1 Gross anatomy There are two classic German references (Tietze, 1921 and Blechschmidt, 1934) on the anatomy of the heel pad that were translated into English [1,2]. Tietze employed tissue slices and X-rays taken with a soft tissue technique to study the sole of the foot [1]. The author depicted the heel pad as a body of fat that surrounds the calcaneus like a hood and the connective tissue dispersed within the fat was described as mechanical in nature. The individual fat cells and globules were recognized to be sealed off by the connective tissue membranes. Blechschmidt made serial sections of embryonic and adult calcaneal heel pads in three planes (Fig. 8.1) [2]. Extensive descriptions of the patterns and differences in the septa were offered for the fifth, sixth, and seventh embryonic months; by the seventh month, the structure of the heel pad was almost complete. The septa were arranged in a “turbine-like shape” which was referred to as the Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00045-7 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 8.1 Coronal frozen section through the calcaneus of an adult. 60 mm from left to right. The finer interior structure are elastic septa; these chambers are filled with adipocytes [2].

“whorl.” The adult heel pad was found to have septa below the calcaneus that change from a transverse into an oblique presentation and then rearrange into the whorl. Kuhns provided a qualitative description of the normal heel pad as well as heel pads from people who were older, obese, or had suffered from heel pain or some trauma [3]. The author noted changes with elderly subjects (the calcanei of elderly subjects were rough and bony, evidence of a loss of elasticity of the heel pad) and people with heel pain or obesity or trauma also demonstrated bony proliferation. Kuhns noted that the elastic septa in elderly patient’s heels were often fractured and distorted.

8.2.2 Histological or biochemical Kimani provided a detailed discussion of the organization of the connective tissue in the sole of the foot with particular emphasis on elastic fibers and collagen [4]. The sole of the foot was described as having five layers: the epidermis, papillary dermis, reticular dermis, superficial subcutaneous stratum, and deep subcutaneous stratum. Thick fibrous strands bound the dermis to the subcutaneous tissue. The strands enclosed compartments of adipose tissue and then attached to a septum (likely the panniculus carnosus described below) that divided the subcutaneous tissue into the superficial and deep layers. Jahss et al. performed detailed anatomical, histological, and histochemical studies of normal and abnormal heel pads on cadaveric feet [5]. Normal feet have globules of fat surrounded by septa that are attached superficially to the skin. In dysvascular feet, there was less fat and the elastic fibers were more numerous, thicker, and fragmented. In both normal and dysvascular feet there were no interconnections between the “honeycombed” units and the adipose tissue did not flow between compartments. Despite the frayed appearance of the septa in dysvascular feet, their mechanical integrity was not completely lost. Buschmann et al. employed light and electron microscopy to perform a histologic examination of feet with normal and atrophic heel pads [6]. Normal heel pads were found to have a thick epidermis attached to the underlying papillary dermis. The papillary dermis and the reticular dermis consisted of intertwined elastic and collagen fibers. Beneath the dermis was the superficial subcutaneous layer of the fat pad, which was separated from the deep layer by a horizontal septum (again, likely the panniculus carnosus). Within the subcutaneous layers were larger vertical septa that were divided by smaller elastic septa. A histomorphometric analysis by the same group demonstrated differences between normal and atrophied heel pads [6]. In the superficial subcutaneous tissue, the adipocytes in normal heels were 25% larger in mean cell area and 10% larger in mean maximum diameter than in atrophic heel pads. In the deep subcutaneous layer, the difference was 45% in area and 25% in diameter. Both differences were significant. The average septal widths also varied between the types of heel pads, with the septa in atrophied heels being significantly thicker (145 vs 90 µm). Cichowitz et al. conducted several dissections, histological analyses, and arterial and venous studies of the foot [7]. As part of this work, they noted that the panniculus carnosus muscle was observed as a layer in the subcutaneous tissue. The authors suggested that this metabolically active tissue layer is involved in ulcer development. Using an invasive technique, Waldecker and Lehr [8] examined biopsied metatarsal fat pad tissue but found no difference in adipocyte size between diabetic and nondiabetic tissue. As such, systematic atrophy of the metatarsal fat pad was not observed.

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Comparing tissue samples from older diabetic and older nondiabetic specimens, Wang et al. conducted histomorphological and biochemical analyses on plantar soft tissue from six locations beneath the foot (hallux, first, third and fifth metatarsal heads, lateral midfoot and calcaneus) [9]. The main findings were that diabetic tissue has increased septal wall thickness (supporting Buschmann’s earlier work) and increased elastic content. However, unlike Buschmann, they found no change in the adipocyte area (Fig. 8.2).

8.2.3 Medical imaging of tissue thickness Perhaps, the most important mechanical characteristic of the plantar soft tissue is the thickness, which has been studied since at least the mid-1960s. While not comprehensive, we will review some of the more important studies that included plantar tissue thickness as a primary measure (as opposed to a study on the mechanical properties that also reported thickness) for both healthy (Table 8.1) and pathologic tissue (Table 8.2). Steinbach and Russell employed lateral radiographs to study the thickness (unloaded thickness from the bone to “ground” or air [i.e., skin plus fat] unless otherwise stated) of heel pads in 103 normal subjects as well as 29 subjects suffering from acromegaly [10]. The thickness of normal heel pads ranged from 13 to 21 mm with a mean of 17.8 mm, while the heel pads of subjects with acromegaly ranged from 17 to 34 mm with an average of 25.6 mm. Kho et al. also explored the effects of acromegaly using lateral radiographs [11]. They examined 52 control subjects (thickness 5 18.6 mm) and 79 pathologic subjects (thickness 5 27.0 mm).

FIGURE 8.2 Elastic septa in (A) normal (e.g., nondiabetic) and (B) diabetic plantar soft tissue stained with modified Hart’s. Elastic fibers (black) are frayed and fragmented in the diabetic tissue, resulting in thicker elastic septa. Scale bar represents 50 µm [9].

TABLE 8.1 The average unloaded thickness of normal heel pads as described in the literature. Author

Year

Method

Number of heels

Average thickness (mm)

Steinbach

1964

Radiography

206

17.8

Kho

1970

Radiograph

104

18.6

Gooding

1985

Ultrasonography

20

16.6

Gooding

1986

Ultrasonography

48

18.6

Prichasuk

1994

Radiography

800

18.7

Prichasuk

1994

Radiography

400

18.8

Silver

1994

Ultrasonography

21

16.1

Uzel

2006

Ultrasonography

220

18.3

Campanelli

2011

CT

9 (4/5)

16/13a

Hall

2014

Ultrasonography

39

13.8

Lo´pez-Lo´pez

2019

Ultrasonography

190

10.4

Belhan

2019

Ultrasonography

50

19.9

a

Male/female isolated fat thickness.

a

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TABLE 8.2 The average unloaded thickness of abnormal heel pads as described in the literature. Author

Year

Method

Malady

Number of heels

Average thickness (mm)

Steinbach

1964

Radiograph

Acromegaly

58

25.6

Kho

1970

Radiograph

Acromegaly

138

27.0

Gooding

1985

Ultrasonography

Diabetes

76

17.8

Gooding

1986

Ultrasonography

Diabetes without ulcers

76

17.3

Gooding

1986

Ultrasonography

Diabetes with ulcers

22

15.8

Prichasuk

1994

Radiograph

Heel pain

70

20.7

Silver

1994

Ultrasonography

Calc. fract.

21

17.4

Lo´pez-Lo´pez

2019

Ultrasonography

Heel pain

185

7.2

Belhan

2019

Ultrasonography

Heel pain

50

19.5

FIGURE 8.3 Drawing of a lateral radiograph of the foot demonstrating heel pad thickness (A). ab, the skin line; cd, the lowest part of the plantar tuberosity of the calcaneus [14].

Gooding et al. examined both the heel pads of healthy (n 5 10) and diabetic (n 5 38) patients without ulcers [12]. Ultrasonography was used to determine that the healthy subjects had an average heel pad thickness of 16.6 6 0.3 mm, while diabetic subjects had an average thickness of 17.8 6 0.3 mm. The control population had an average age of 28 years while the diabetic population averaged 62 years, so differences in thickness could be due to nothing more than age. A second paper by Gooding et al. examined the thickness of the heel pad for normal subjects (n 5 24), diabetic subjects (n 5 38), and diabetic subjects with foot ulcers (n 5 11) [13]. The normal subjects had the thickest heel pads (18.6 mm), significantly larger than the diabetic subjects without ulcers (17.3 mm), which was, in turn, significantly larger than the diabetic subjects with ulcers (15.8 mm). The results are in contrast to the author’s earlier results; however, the three groups in this study were much closer in age (means of 51 6 3.8, 62 6 1.2 and 60 6 2.3 years respectively) and that probably explains the discrepancy. Additionally, the authors used the same methodology to study the thickness of the soft tissue beneath the five metatarsal heads. The soft tissue thicknesses for the normal subjects, from the first to the fifth metatarsal, were found to be: 12.9, 14.2, 13.6, 12.9 and 11.5 mm, respectively. Prichasuk et al. examined the loaded and unloaded thickness of the heels of 400 normal subjects (Fig. 8.3) [14]. The subject population was subdivided into four groups: men aged 20 35, men aged 40 60, women aged 20 35, and women aged 40 60. Unloaded heel pads ranged from 12 to 28 mm with an average of 18.7 mm. The authors determined that heel pads were significantly thicker in male subjects and in older subjects. The elasticity of the heel pad was not dependent on gender but rather on age and weight. Older or heavier subjects had heel pads that were significantly less elastic (i.e., they were stiffer). The lead author in the preceding article also studied 70 subjects with heel pain as well as an additional 200 normal subjects [15]. Prichasuk determined that subjects with heel pain had thicker and stiffer heel pads. Similar to the first

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139

study, males had thicker heel pads but no difference in elasticity when compared to females while older subjects had thicker and stiffer heel pads than younger subjects. Silver et al. examined the heel pad thickness of 21 patients with unilateral fractures of the calcaneus using ultrasound [16]. The uninjured side served as a control. The fractured side was found to have a significant increase in thickness (average 17.4 mm, range 12.0 23.9 mm) compared to the control side (average 16.1 mm, range 11.3 19.8 mm). The effect of athletic activity on heel pad thickness was explored by Uzel et al. [17]. Young adults were considered in three groups: sedentary, ,7 hour activity per week, and .7 hours activity per week, but there was no difference in heel pad thickness (average 18.3 mm, range 14.0 to 24.5 mm across all groups). Hall et al. studied a population of asymptomatic long-distance runners [18], who were found to have thinner heel pads (13.8 mm) that were on the edge of what is considered normal in the literature. Campanelli et al. performed of 3D morphological computed tomography (CT)-based study of the anatomy of 9 young adult heel pads [19]. From the 3D segmented CT scan volume, a series of 1D and 2D measurements were made. The heel pads had a crest on the anterior dorsal surface, flanges medially, laterally, and posteriorly, and a portion that covered the Achilles tendon insertion. Male heel pads were thicker than female heel pads. Lo´pez-Lo´pez examined 375 patients who were categorized into two groups: heel pain and controls [20]. They used ultrasonography to measure heel pad thickness. The subjects with heel pain had significantly thinner heel pads (7.2 mm vs 10.4 mm). Of note, the authors did not explain why their methods produced much thinner heel pads than other comparable studies. Another group examined the plantar fat pad thickness (both heel and first metatarsal) of patients with unilateral heel pad compared to their contralateral pain free limb [21]. Subjects with heel pain had small (0.5 mm) but significantly thinner heel pads (19.4 vs 19.9 mm) with similar first metatarsal thicknesses (6.8 and 6.8 mm). In addition to the heel pad thickness work, some groups have also examined the thickness of the forefoot plantar soft tissue using ultrasound (Table 8.3). The work by Gooding et al. studied the five metatarsal heads for controls, diabetic subjects without ulcers, and diabetic subjects with ulcers [13]. Bygrave and Betts, studied the thickness of the plantar soft tissue at the five metatarsal heads and the two sesamoids while the foot was loaded and unloaded; only the unloaded metatarsal data are reported here [22]. Cavanagh employed a specially mounted ultrasound transducer to capture the second metatarsal thickness during walking [23]; they also capture the unloaded thickness. Wang et al. develop a novel ultrasonic device to measure forefoot tissue thickness beneath all the metatarsal heads during quiet stance after measuring the unloaded thickness [24]. Mickle et al. explored the effect of both hallux valgus and lesser toe deformities on plantar soft tissue thickness at the first and fifth metatarsal head, as well as the heel [25]. They found decrease first metatarsal thickness with hallux valgus and decrease fifth metatarsal thickness with lesser toe deformities. The last anatomical aspect of the plantar soft tissue that will be discussed is the concept of micro and macrochambers. As detailed in the literature [2,6,26], the plantar skin and soft tissue can be considered a layered structure

TABLE 8.3 The average unloaded thickness of the metatarsal heads as described in the literature. Author

Group

Number for each metatarsal head

First-met mean (error)a mm

Secondmet mean (error)a mm

Third-met mean (error)a mm

Fourth-met mean (error)a mm

Fifth-met mean (error)a mm

Gooding

Control

48/48/48/46/47

12.92 (0.42)

14.2 (0.3)

13.6 (0.3)

12.9 (0.3)

11.5 (0.3)

Diabetic without ulcers

75/76/76/76/76

11.6 (0.3)

12.7 (0.3)

13.1 (0.24)

12.1 (0.3)

10.7 (0.2)

Diabetic with ulcers

21/22/22/19/20

10.7 (0.6)

12.5 (0.7)

12.3 (0.5)

11.1 (0.5)

10.7 (0.4)

Bygrave and Betts

Control

24 each

12.8 (3.3)

10.47 (2.4)

8.58 (2.2)

7.77 (1.8)

7.77 (1.8)

Cavanagh

Control

5 (second only)

Wang

Control

50

13.7 (2.1)

11.8 (1.8)

11.1 (1.7)

10.3 (1.2)

Belhan

Heel pain/ control

50/50

6.8/6.8

15.2 12.6 (1.9)

Error 5 standard error for Gooding; 2 standard deviations for Bygrave and Betts; 1 standard deviation (assumed) for Wang.

a

140

PART | 2 Function

consisting of the epidermis and dermis (skin), followed by the superficial adipose layer, a muscle layer, and the deep adipose layer. Ultrasound has been used to demonstrate that the superficial layer and deep layer consist of microchambers and macro-chambers [27]; this distinction between micro-chamber and macro-chamber has implications for plantar soft tissue mechanics [28 30]. Recently, investigators have used machine learning techniques to help quantify differences between the superficial and deep layers [31].

8.3

Biomechanical function

The plantar soft tissue serves several important functions in the body. As summarized by Rome pertaining to the heel pad, but extending to the rest of the plantar soft tissue, these include shock reduction, energy absorption, and load distribution [32]. Shock reduction refers to the deceleration of an effective mass over a certain distance; this leads to a reduction in peak force or loading rate. By considering the peak force at heel strike, the vertical velocity at heel strike, and mechanical properties of the heel pad, the time to reach peak force and the effective mass of the heel pad were estimated [33]. The implication is that a stiffer heel pad would lead to great shock imparted to the body. Energy absorption, referred to by Rome as shock absorption, is related to the deformation of the plantar soft tissue to dissipate energy. Jahss demonstrated that fluid (fat) does not flow between septal chambers [34]. Thus as the plantar tissue is loaded in compression, it bulges and stresses the internal reinforcement structures, i.e., the elastic septa. If these septa become damaged or frayed, as in the case of diabetes, they are more likely to deform and less able to resist compression [5]. The tissues are also less likely to recover once the load is removed. These changes can affect the energy dissipation of the plantar soft tissue [2]. Finally, the heel pad alters the distribution of plantar pressure; normal heel pads have a broad region of higher pressure, while atrophic heel pads have a narrow peak pressure zone [34]. More recently, Bus et al. demonstrated that diabetic neuropathic subjects had a strong inverse correlation between an magnetic resonance imaging (MRI)-determined heel pad fat fraction and peak heel plantar pressure [35], further indicating that the plantar fat is related to the normal distribution of vertical force.

8.4

Mechanical properties

The mechanical properties of plantar soft tissue have been explored from a structural (geometry dependent) and a material (geometry independent) standpoint in both living and cadaveric subjects.

8.4.1 Structural in vivo testing Cavanagh et al. employed a pendulum impact tester to conduct experiments on 10 subjects [36]. The protocol involved two impact velocities (1.03 and 1.44 m/s) and provided the peak acceleration (20.8 and 36.3 g), peak force (338 and 676 N), maximum deformation (8.8 and 10.9 mm), and final stiffness (139 kN/m) as well as the energy loss (85% and 90%) (Table 8.4). Using a similar set up, Gordon Valiant conducted pendulum impact testing on the heel pads of 12 runners and 12 non-runners [37,38]. Findings were consistent with Cavanagh’s work (they were collaborators), with peak force (223 and 437 N), maximum deformation (8.5 and 9.8 mm), final stiffness (106 kN/m), and energy loss (84% and 99%) very similar. Kinoshita et al. used a vertical impact tester to provide normative data on the heel pads of 16 adults and seven children [39]. The study incorporated two drop heights (30 and 50 mm) which resulted in peak accelerations of 8.7 and 11.6 g respectively. For these drop heights, the average maximum deformations (9.5 and 11.3 mm) and energy losses (77.4% and 78.8%) were reported. The children were found to have larger peak accelerations and maximum deformations as well as smaller percent energy losses. The author’s employed the same apparatus to look at 10 young adults (age 17 30 years) as well as two groups of active elderly individuals (age 60 67 years and 71 86 years) [40]. For two drop heights (20 and 50 mm), the young adults had peak accelerations of 6.8 and 10.8 g, maximum deformations of 8.6 and 11.2 mm, and energy losses of 76.5% and 78.8%, respectively. At the lower drop height, both of the elderly groups had significantly less energy absorption than the young adults, but there was no difference in peak acceleration or maximum deformation.

8.4.2 Structural ex vivo testing Bennett and Ker studied the mechanical properties of the heel pads of 11 cadaveric limbs, and all of which were amputated due to vascular problems (Table 8.5) [41]. The ex vivo specimens, some attached to the calcaneus and some not,

TABLE 8.4 Test parameters and results of several in vivo impact testing experiments on the subcalcaneal tissue. First author

Year

n

Mass (kg)

Velocity (m/s)

Drop height (mm)

Energy (J)

Peak acceleration (g)

Maximum deformation (mm)

Peak force (N)

Initial stiffness (kN/m)

Final stiffness (kN/m)

Energy loss (%)

Cavanagh

1984

10

1.9

1.03, 1.44

54,* 106*

1.0,* 1.97*

20.8, 36.3

8.8, 10.9

338, 676

19

138

85, 90

Valiant

1984

24

1.92

0.8, 1.0, 1.2

33,* 51,*, 73*

0.61,0.96, 1.38

na

8.5, 9.8

223, 437

7.91

106

84, 99

Kinoshita

1993

16

5

0.72, 0.93

30, 50

1.3, 2.16

8.7, 11.6

9.5, 11.3

427, 569

25*, 25*

100*, 175*

77.4, 78.8

Kinoshita

1996

10

5

0.57, 0.94

20, 50

0.81, 2.2

6.8, 10.8

8.6, 11.2

333, 531

37.5*, 37.5*

125*, 125*

76.5, 78.1

Kinoshita

1996

10

5

0.57, 0.94

20, 50

0.81, 2.2

7.0, 11.9

8.5, 10.3

344, 584

25*, 25*

175*, 175*

73.8, 73.6

Kinoshita

1996

10

5

0.57, 0.94

20, 50

0.81, 2.2

7.6, 12.6

8.6, 10.4

372, 616

25*, 25*

175*, 175*

72.6, 72.6

Cavanagh and Valiant employed a pendulum impact tester while Kinoshita used a vertical impact tester. Note that Kinoshita’s 1996 work was presented in three rows, representing young adults (age 17 30 years), seniors (age 60 67 years) and older seniors (age 71 86 years). Estimated parameters are signified with an *.

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PART | 2 Function

TABLE 8.5 Test parameters and results of several ex vivo impact testing experiments on the subcalcaneal tissue. First author

Year

n

Frequency (Hz)

Peak force (N)

Final stiffness (kN/m)

Maximum deformation (mm)

Energy loss (%)

Bennett

1990

11

0.1 70

1200

1160

2.1, 4.7

28.6, 32.3

Ker

1996

3

5.5

1400

950

3.67

32.8

Aerts

1995

5

1.1 11

2000

900

4 5

46.5 65.5

Bennett reported two displacements, for forces of one and two times body weight, and two energy loss values, for the heel pad attached and unattached to the calcaneus. The range of deformation from Aerts et al. was controlled by the experimental protocol, while the two energy loss values represent the nth loop and the first loop.

were tested with an Instron materials testing machine up to applied loads of 2 kN. A range of frequencies (0.1 70 Hz) and of temperatures (0 C to 37 C) were employed. The results demonstrated no significant difference in stiffness or energy dissipation over the range of frequencies tested. A small, but significant difference in energy dissipation was seen from 0 C to 37 C, but no differences were found when a specimen was tested, frozen and retested on a later date. Ker performed mechanical tests on three ex vivo heel pads from feet removed due to irreparable vascular disease [42]. The author was interested in stimulating the tissue in a manner similar to gait, (i.e., with delays in between loadings). Ker developed a multiple regression equation whereby energy loss was the dependent variable while delay time and peak force were the independent variables. The author did demonstrate that the energy loss of the heel pads increased with the amount of rest time between tests, and he reported averages for energy loss (32.8%) and stiffness (0.95 kN/m). Aerts et al., noting discrepancies in the mechanical testing protocols, modified an Instron ex vivo oscillatory protocol to impart an impact half way through the cycle (i.e., at maximum velocity) and performed pendulum impact tests on the mechanically grounded subcalcaneal tissue of ex vivo (as opposed to in vivo) specimens [43]. Results were similar for both tests and included a 46.5% 65.5% energy absorption, a 4 to 5 mm deformation, and a stiffness of 9.0 3 105 N/m. Contrasting the impacting testing with the cadaveric work (Tables 8.4 and 8.5) demonstrates that in vivo results were less stiff, with increased deformation and energy loss. The modified protocols by Aerts et al. were in between the in vivo and ex vivo tests, but 4 of the 5 feet were from specimens between the ages of 50 and 80 that had been removed due to vascular disease and tested at room temperature. This is potentially a concerns, since age, diabetic status, and temperature can affect the mechanical properties of the plantar soft tissue [44].

8.4.3 Material ex vivo testing Several studies have quantified the compressive material properties of the subcalcaneal tissue. Miller-Young et al. demonstrated that the plantar soft tissue is isotropic; they developed an analytical model for the soft tissue that incorporates hyperelastic and viscoelastic characteristics. While this research was the first compressive material property data published on the plantar soft tissue, they only tested heel pad tissue from older feet (age 61 99) of unknown vascular condition and experiments were conducted at room temperature. In a series of tests, Ledoux et al. explored the compressive and shear material properties of the plantar soft tissue, while addressing testing conditions related to age, diabetic status, and temperature. A detail review of their work is beyond the scope of this chapter, but it included the compressive material properties, with an initial [44] and refined [45] testing protocol (Fig. 8.4), and shear material properties [46]. They also developed quasilinear viscoelastic models for structural punch testing [47] and materials testing [48], and hyperelastic inverse finite element derived models [49]. Finally, they investigated the effect of target strain error [50] and of prior testing on the compressive material properties [51]. The fundamental finding of this work was that the modulus of healthy plantar tissue is approximately 600 kPa with energy absorption of roughly 70% [45].

8.4.4 Ultrasound Ultrasound has been used extensively to explore the plantar soft tissue. Initially, the measurement systems consisted of pen-sized devices with an ultrasound transducer in series with a load cell [52]. This device was used to compare 4 older

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143

FIGURE 8.4 (A) Specimen locations at the hallux (ha), first, third, and fifth metatarsal heads (m1, m3, and m5), lateral midfoot (la), and calcaneus (ca) as well as (B) a typical plantar tissue specimen before skin removal. (C) Experimental set-up showing specimen in environmental chamber between sandpaper covered platens and (D) after sealing to maintain in vivo conditions of near 100% humidity and 35 C. [45].

diabetic neuropathic patients with 4 young healthy individuals; the former were significantly thinner and stiffer [53]. Wang et al. used an ultrasound probe in series with a push-pull scale to load the forefoot and estimate the elastic modulus and energy dissipation ratio, which were found to vary with bodyweight in normal subjects [54]. To obtain plantar soft tissue properties during functional loading, one group embedded an ultrasound sensor in shoes to obtain dynamic soft tissue deformation in 16 normal subjects while exploring the effect of a heel cup [55]. More recently, shear wave elastography has been employed to quantify the stiffness of the plantar soft tissue, demonstrating a pattern of decreased stiffness from superficial to deep [56] and to explore healthy soft tissue properties [57]. As discussed elsewhere in this chapter, many groups used ultrasound systems to study the effects of aging or diabetes on mechanical characteristics of the plantar soft tissue.

8.4.5 Other in vivo techniques The inverse finite element method has been employed to estimate the mechanical properties of the plantar soft tissue. Initial 2D work demonstrated no differences between diabetic and nondiabetic tissue, which indicated the importance of patient specific modeling [58]. This was expanded to develop 3D specimen-specific models, suggesting the framework could be expanded to predict how the heel pad would interact with the surroundings [59]. Others have incorporated inverse finite element models with ultrasound scanning techniques to develop patient specific models [60,61]. (Inverse finite element models were also generated for cadaveric plantar tissue [49].) MRI has been used to study the plantar soft tissue by several groups. This has included: indenting while MRI scanning to measure shear and elastic moduli [62]; using hydraulic cylinders to applying normal and shear loads while conducting MRI scans to quantify internal tissue deformation [63]; and conducting hydraulic-driven, cyclic loading while employing gatedMRI scanning techniques to quantify plantar soft tissue properties [64,65]. Additionally, using a hand held force gage, an indentor, and a linear variable differential transformer, heel pad stiffness was found to be increased in subjects with heel pain compared to those without [66]. Another group developed a load cell in series with a 3D measurement device to conduct tissue indentation studies [67]; they used this device to demonstrate that diabetic plantar soft tissue was stiffer than control tissue [68]. Single plane fluoroscopy has also been used by several research groups to determine in vivo stiffness [69] or in vivo modulus [70,71].

8.5

Effect of aging

Biological tissues often demonstrate changes in material and structure characteristics as the tissue grows and develops, and then again as the tissue ages past maturity. For the purposes of this discussion, we are only interested in changes that are part of the aging process. There have been many studies on this issue related to the plantar soft tissue; a brief review is presented hereafter. Using a linear impact tester and studying the heel pads of living subjects, Kinoshita et al. found that with age comes diminished energy absorption [40]. An ultrasound transducer in series with a push-pull scale was used to compare subjects younger than 40 to subjects older than 60 [72]. The older tissue was thinner, less elastic, and less able to absorb energy. A custom frame allowed for the manual application of force to the metatarsal heads of young (19 35 years) and older (42 72 years) subjects via an ultrasound probe in series with a load cell [73]. In general, the older tissue was stiffer, allowed for less strain, and absorbed more energy. In another study based on an ultrasound probe and load cell in series, Kwan et al. found no change in thickness, but a significant increase in stiffness across multiple locations on the plantar soft tissue (hallux, first, third, and fifth metatarsal heads, and the calcaneus) [74].

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PART | 2 Function

A motorized indentor in series with a load cell was used to quantify second metatarsal stiffness for young (mean age 22 years) and old (mean age 67 years) subjects. Across varying metatarsophalangeal angles (0, 20, and 40 degrees), the older specimens were significantly stiffer [75]. Shear wave elastography has been used to study the deep and superficial layer differences between younger (26 years) and older (67 years) subjects [29]; older subjects had increased overall and superficial stiffness, with a slightly decreased deep layer stiffness. However, another shear wave elastography study found that the superficial layer stiffness decreased with age when comparing subjects in their 20s with subjects in their 60s [30]. In summary, despite some conflicting results, most research has demonstrated that older plantar soft tissue is thinner, less elastic, and stiffer, with diminished energy absorption.

8.6

Diabetic plantar soft tissue

Common methods for investigating plantar soft tissue include cadaveric testing, finite element models, MRI, and ultrasound. Cadaveric investigations have found increased modulus in compressive (Fig. 8.5) [45] and shear loading [46], increased total plantar soft tissue thickness [76], increased dermal thickness [26], and thicker, frayed septal walls [5,6,26] in diabetic specimens compared to nondiabetic specimens. Finite element models have been used to design therapeutic footwear [77,78], estimate effects of altered tissue mechanics on plantar pressure [79], and to calculate plantar soft tissue properties [60,61]. MRI has also been used to study diabetic plantar soft tissue mechanics in vivo [65]. Ultrasound has been used extensively for in vivo measurement of soft tissue deformation and this technique has been used to explore various aspects of the effect of diabetes on the plantar soft tissue. Studies have used ultrasound probes in series with load cells to explore differences in deep vs superficial plantar soft tissue layers between diabetic and nondiabetic subjects [28], and measured an increase in stiffness with diabetic specimens [76,80].

8.6.1 Other pathologies associated with the plantar soft tissue Caused by over production of growth hormone, acromegaly leads to abnormal growth of the hands, feet, and face. Increased heel pad thickness has been used to diagnose this disease, with an acromegalic group having a heel pad thickness of 25.6 mm compared to 17.8 mm for a control group [10]. Kho et al. found very similar results, with male/female patients with acromegaly having heel pad thicknesses of 28.0 and 26.0 mm respectively, while age-matched male/ female controls were 18.6 and 18.5 mm thick [11]. Long term Dilantin treatment, used to control seizures, has been associated with increased heel pad thickness; patients on Dilantin for 10 years or more had a mean heel pad thickness of 20.6 mm, compared to 16.6 mm for the control group [81]. However, despite hypotheses to the contrary [82], heel pad damage (edema, fibrosis, fatty release, loss of thickness, or septa damage) was not found in a prospective study of 22 calcaneal fractures [83]; in fact, others have found a significant increase in heel pad thickness post calcaneal fracture compared to the contralateral limb [16]. Finally, contradictory information has been found for changes related to heel pain. Early work found that compared to a control group, painful heel pads were thicker [15], but a recent study has demonstrated that heel pads with ipsilateral heel pain are thinner than that of the contralateral pain free side [21]. These differences are likely due to the demographic differences between groups in the earlier study; heel pain subjects were older and heavier.

FIGURE 8.5 Modulus as a function of (A) frequency (across all locations) and (B) location (across all frequencies) where N, nondiabetic, D, diabetic, ha 5 hallux, m1, m3, and m5 5 first, third, and fifth metatarsals, la 5 lateral midfoot, and ca 5 calcaneus [45].

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Areas of future biomechanical research

Our understanding of the mechanics of the plantar soft tissue has increased profoundly as measuring tissue thickness via X-ray was the state-of-the-art. Exciting advancements have been made with shear wave elastography, MRI scanning, and computational modeling that have the field on the cusp of conducting patient-specific analyses to further elucidate changes due to aging or the diabetes disease process. One gap in our knowledge is a mechanistic understanding of how biochemical and histological characteristics are associated with mechanical properties. Combining this understanding with the available imaging and modeling techniques will allow for advancements in treatment strategies, such as custom orthoses, that can mitigate diabetic foot complications and ultimately reduce ulceration rates.

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Tietze A. Concerning the architectural structure of the connective tissue in the human sole. Foot Ankle 1982;2(5):252 9. Blechschmidt E. The structure of the calcaneal padding. Foot Ankle 1982;2(5):260 83. Kuhns JG. Changes in elastic adipose tissue. J Bone Joint Surg 1949;31-A(3):541 7. Kimani JK. The structural and functional organization of the connective tissue in the human foot with reference to the histomorphology of the elastic fibre system. Acta Morphol Neerl Scand 1984;22(4):313 23. Jahss MH, Michelson JD, Desai P, Kaye R, Kummer F, Buschman W, et al. Investigations into the fat pads of the sole of the foot: anatomy and histology. Foot Ankle 1992;13(5):233 42. Buschmann WR, Jahss MH, Kummer F, Desai P, Gee RO, Ricci JL. Histology and histomorphometric analysis of the normal and atrophic heel fat pad. Foot Ankle Int 1995;16(5):254 8. Cichowitz A, Pan WR, Ashton M. The heel: anatomy, blood supply, and the pathophysiology of pressure ulcers. Ann Plast Surg 2009;62 (4):423 9. Waldecker U, Lehr HA. Is there histomorphological evidence of plantar metatarsal fat pad atrophy in patients with diabetes? J Foot Ankle Surg 2009;48(6):648 52. Wang YN, Lee K, Shofer JB, Ledoux WR. Histomorphological and biochemical properties of plantar soft tissue in diabetes. Foot (Edinb) 2017;33:1 6. Steinbach HL, Russell W. Measurement of the heel pad as an aid to diagnosis of acromegaly. Radiology 1964;82:418 22. Kho KM, Wright AD, Doyle FH. Heel pad thickness in acromegaly. Br J Radiol 1970;43(506):119 25. Gooding GA, Stress RM, Graf PM, Grunfeld C. Heel pad thickness: determination by high-resolution ultrasonography. J Ultrasound Med 1985;4(4):173 4. Gooding GA, Stess RM, Graf PM, Moss KM, Louie KS, Grunfeld C. Sonography of the sole of the foot. Evidence for loss of foot pad thickness in diabetes and its relationship to ulceration of the foot. Invest Radiol 1986;21(1):45 8. Prichasuk S, Mulpruek P, Siriwongpairat P. The heel-pad compressibility. Clin Orthop Relat Res 1994;300:197 200. Prichasuk S. The heel pad in plantar heel pain. J Bone Joint Surg Br 1994;76(1):140 2. Silver DA, Kerr PS, Andrews HS, Atkins RM. Heel pad thickness following calcaneal fractures: ultrasound findings. Injury 1994;25(1): 39 40. Uzel M, Cetinus E, Ekerbicer HC, Karaoguz A. Heel pad thickness and athletic activity in healthy young adults: a sonographic study. J Clin Ultrasound 2006;34(5):231 6. Hall MM, Finnoff JT, Sayeed YA, Smith J. Sonographic evaluation of the plantar heel in asymptomatic endurance runners. J Ultrasound Med 2015;34(10):1861 71. Campanelli V, Fantini M, Faccioli N, Cangemi A, Pozzo A, Sbarbati A. Three-dimensional morphology of heel fat pad: an in vivo computed tomography study. J Anat 2011;. Lopez-Lopez D, Becerro-de-Bengoa-Vallejo R, Losa-Iglesias ME, Soriano-Medrano A, Palomo-Lopez P, Morales-Ponce A, et al. Relationship between decreased subcalcaneal fat pad thickness and plantar heel pain. A case control study. Pain Physician 2019;22(1):109 16. Belhan O, Kaya M, Gurger M. The thickness of heel fat-pad in patients with plantar fasciitis. Acta Orthop Traumatol Turc 2019;53(6):463 7. Bygrave CJ, Betts RP. The plantar tissue thickness in the foot: a new ultrasound technique for loadbearing measurements and a metatarsal head depth study. The Foot 1992;2:71 8. Cavanagh PR. Plantar soft tissue thickness during ground contact in walking. J Biomech 1999;32(6):623 8. Wang TG, Hsiao TY, Wang TM, Shau YW, Wang CL. Measurement of vertical alignment of metatarsal heads using a novel ultrasonographic device. Ultrasound Med Biol 2003;29(3):373 7. Mickle KJ, Munro BJ, Lord SR, Menz HB, Steele JR. Soft tissue thickness under the metatarsal heads is reduced in older people with toe deformities. J Orthop Res 2011;29(7):1042 6. Wang YN, Lee K, Ledoux WR. Histomorphological evaluation of diabetic and non-diabetic plantar soft tissue. Foot Ankle Int 2011;32 (8):802 10. Hsu CC, Tsai WC, Wang CL, Pao SH, Shau YW, Chuan YS. Microchambers and macrochambers in heel pads: are they functionally different? J Appl Physiol 2007;102(6):2227 31.

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[28] Hsu CC, Tsai WC, Hsiao TY, Tseng FY, Shau YW, Wang CL, et al. Diabetic effects on microchambers and macrochambers tissue properties in human heel pads. Clin Biomech (Bristol, Avon) 2009;24(8):682 6. [29] Wu CH, Lin CY, Hsiao MY, Cheng YH, Chen WS, Wang TG. Altered stiffness of microchamber and macrochamber layers in the aged heel pad: shear wave ultrasound elastography evaluation. J Formos Med Assoc 2018;117(5):434 9. [30] Mo F, Li J, Yang Z, Zhou S, Behr M. In vivo measurement of plantar tissue characteristics and its indication for foot modeling. Ann Biomed Eng 2019;47(12):2356 71. [31] Brady L, Wang Y-N, Rombokas E, Ledoux WR. Adipose chamber size and automated feature extraction in diabetic plantar soft tissue. In: Proceedings of the 45th annual meeting of the American Society of Biomechanics. Atlanta, GA (virtual); 2021. [32] Rome K. Mechanical properties of the heel pad: current theory and review of the literature. The Foot 1998;8:179 85. [33] Ker RF, Bennett MB, Alexander RM, Kester RC. Foot strike and the properties of the human heel pad. Proc Inst Mech Eng 1989;203 (4):191 6. [34] Jahss MH, Kummer F, Michelson JD. Investigations into the fat pads of the sole of the foot: heel pressure studies. Foot Ankle 1992; 13(5):227 32. [35] Bus SA, Akkerman EM, Maas M. Changes in sub-calcaneal fat pad composition and their association with dynamic plantar foot pressure in people with diabetic neuropathy. Clin Biomech (Bristol, Avon) 2021;88:105441. [36] Cavanagh PR, Valiant GA, Misevich KW. Biological aspects of modeling shoe/foot interaction during running. In: Fredericks EC, editor. Sports shoes and playing surfaces: biomechanical properties. Champaign, IL: Human Kinetics Publishers, Inc; 1984. p. 24 46. [37] Valiant GA. Transmission and attenuation of heelstrike accelerations. In: Cavanagh P, editor. Biomechanics of distance running. Champaign, IL: Human Kinetics Books, Inc.; 1990. p. 225 47. [38] Valiant GA, Cavanagh PR. An in vivo determination of the mechanical characteristics of the human heel pad. J Biomech 1985;18:242. [39] Kinoshita H, Ogawa T, Kuzuhara K, Ikuta K. In vivo examination of the dynamic properties of the human heel pad. Int J Sports Med 1993; 14(6):312 19. [40] Kinoshita H, Francis PR, Murase T, Kawai S, Ogawa T. The mechanical properties of the heel pad in elderly adults. Eur J Appl Physiol Occup Physiol 1996;73(5):404 9. [41] Bennett MB, Ker RF. The mechanical properties of the human subcalcaneal fat pad in compression. J Anat 1990;171:131 8. [42] Ker RF. The time-dependent mechanical properties of the human heel pad in the context of locomotion. J Exp Biol 1996;199(Pt 7):1501 8. [43] Aerts P, Ker RF, De Clercq D, Ilsley DW, Alexander RM. The mechanical-properties of the human heel pad: a paradox resolved. J Biomech 1995;28(11):1299 308. [44] Ledoux WR, Blevins JJ. The compressive material properties of the plantar soft tissue. J Biomech 2007;40(13):2975 81. [45] Pai S, Ledoux WR. The compressive mechanical properties of diabetic and non-diabetic plantar soft tissue. J Biomech 2010;43(9):1754 60. [46] Pai S, Ledoux WR. The shear mechanical properties of diabetic and non-diabetic plantar soft tissue. J Biomech 2012;45(2):364 70. [47] Ledoux WR, Meaney DF, Hillstrom HJ. A quasi-linear, viscoelastic, structural model of the plantar soft tissue with frequency-sensitive damping properties. J Biomech Eng 2004;126(6):831 7. [48] Pai S, Ledoux WR. The quasi-linear viscoelastic properties of diabetic and non-diabetic plantar soft tissue. Ann Biomed Eng 2011; 39(5):1517 27. [49] Isvilanonda V, Iaquinto JM, Pai S, Mackenzie-Helnwein P, Ledoux WR. Hyperelastic compressive mechanical properties of the subcalcaneal soft tissue: an inverse finite element analysis. J Biomech 2016;. [50] Pai S, Ledoux WR. The effect of target strain error on plantar tissue stress. J Biomech Eng 2010;132(7):071001. [51] Pai S, Vawter PT, Ledoux WR. The effect of prior compression tests on the plantar soft tissue compressive and shear properties. J Biomech Eng 2013;135(9):94501. [52] Zheng YP, Mak AF. An ultrasound indentation system for biomechanical properties assessment of soft tissues in-vivo. IEEE Trans Biomed Eng 1996;43(9):912 18. [53] Zheng YP, Choi YK, Wong K, Chan S, Mak AF. Biomechanical assessment of plantar foot tissue in diabetic patients using an ultrasound indentation system. Ultrasound Med Biol 2000;26(3):451 6. [54] Wang CL, Hsu TC, Shau YW, Shieh JY, Hsu KH. Ultrasonographic measurement of the mechanical properties of the sole under the metatarsal heads. J Orthop Res 1999;17(5):709 13. [55] Telfer S, Woodburn J, Turner DE. Measurement of functional heel pad behaviour in-shoe during gait using orthotic embedded ultrasonography. Gait Posture 2014;39(1):328 32. [56] Lin CY, Chen PY, Shau YW, Tai HC, Wang CL. Spatial-dependent mechanical properties of the heel pad by shear wave elastography. J Biomech 2017;53:191 5. [57] Chatzistergos PE, Behforootan S, Allan D, Naemi R, Chockalingam N. Shear wave elastography can assess the in-vivo nonlinear mechanical behavior of heel-pad. J Biomech 2018;80:144 50. [58] Erdemir A, Viveiros ML, Ulbrecht JS, Cavanagh PR. An inverse finite-element model of heel-pad indentation. J Biomech 2006;39(7):1279 86. [59] Chokhandre S, Halloran JP, van den Bogert AJ, Erdemir A. A three-dimensional inverse finite element analysis of the heel pad. J Biomech Eng 2012;134(3):031002. [60] Chen WM, Lee SJ, Lee PVS. The in vivo plantar soft tissue mechanical property under the metatarsal head: implications of tissues joint-angle dependent response in foot finite element modeling. J Mech Behav Biomed Mater 2014;40:264 74.

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[61] Behforootan S, Chatzistergos PE, Chockalingam N, Naemi R. A clinically applicable non-invasive method to quantitatively assess the viscohyperelastic properties of human heel pad, implications for assessing the risk of mechanical trauma. J Mech Behav Biomed Mater 2017; 68:287 95. [62] Gefen A, Megido-Ravid M, Azariah M, Itzchak Y, Arcan M. Integration of plantar soft tissue stiffness measurements in routine MRI of the diabetic foot. Clin Biomech (Bristol, Avon) 2001;16(10):921 5. [63] Petre M, Erdemir A, Cavanagh PR. An MRI-compatible foot-loading device for assessment of internal tissue deformation. J Biomech 2008; 41(2):470 4. [64] Williams ED, Stebbins MJ, Cavanagh PR, Haynor DR, Chu B, Fassbind MJ, et al. The design and validation of a magnetic resonance imagingcompatible device for obtaining mechanical properties of plantar soft tissue via gated acquisition. Proc Inst Mech Eng H 2015;229(10):732 42. [65] Williams ED, Stebbins MJ, Cavanagh PR, Haynor DR, Chu B, Fassbind MJ, et al. A preliminary study of patient-specific mechanical properties of diabetic and healthy plantar soft tissue from gated magnetic resonance imaging. Proc Inst Mech Eng H 2017;231(7):625 33. [66] Rome K, Webb P, Unsworth A, Haslock I. Heel pad stiffness in runners with plantar heel pain. Clin Biomech 2001;16(10):901 5. [67] Klaesner JW, Commean PK, Hastings MK, Zou D, Mueller MJ. Accuracy and reliability testing of a portable soft tissue indentor. IEEE Trans Neural Syst Rehabil Eng 2001;9(2):232 40. [68] Klaesner JW, Hastings MK, Zou DQ, Lewis C, Mueller MJ. Plantar tissue stiffness in patients with diabetes mellitus and peripheral neuropathy. Arch Phys Med Rehabil 2002;83(12):1796 801. [69] De Clercq D, Aerts P, Kunnen M. The mechanical characteristics of the human heel pad during foot strike in running: an in vivo cineradiographic study. J Biomech 1994;27(10):1213 22. [70] Gefen A, Megido-Ravid M, Itzchak Y. In vivo biomechanical behavior of the human heel pad during the stance phase of gait. J Biomech 2001;34(12):1661 5. [71] Wearing SC, Smeathers JE, Urry SR, Sullivan PM, Yates B, Dubois P. Plantar enthesopathy: thickening of the enthesis is correlated with energy dissipation of the plantar fat pad during walking. Am J Sports Med 2010;38(12):2522 7. [72] Hsu TC, Wang CL, Tsai WC, Kuo JK, Tang FT. Comparison of the mechanical properties of the heel pad between young and elderly adults. Arch Phys Med Rehabil 1998;79(9):1101 4. [73] Hsu CC, Tsai WC, Chen CP, Shau YW, Wang CL, Chen MJ, et al. Effects of aging on the plantar soft tissue properties under the metatarsal heads at different impact velocities. Ultrasound Med Biol 2005;31(10):1423 9. [74] Kwan RL, Zheng YP, Cheing GL. The effect of aging on the biomechanical properties of plantar soft tissues. Clin Biomech (Bristol, Avon) 2010;25(6):601 5. [75] Teoh JC, Shim VP, Lee T. Quantification of plantar soft tissue changes due to aging in various metatarsophalangeal joint angles with realistic tissue deformation. J Biomech 2014;47(12):3043 9. [76] Chao CY, Zheng YP, Cheing GL. Epidermal thickness and biomechanical properties of plantar tissues in diabetic foot. Ultrasound Med Biol 2011;37(7):1029 38. [77] Erdemir A, Saucerman JJ, Lemmon D, Loppnow B, Turso B, Ulbrecht JS, et al. Local plantar pressure relief in therapeutic footwear: design guidelines from finite element models. J Biomech 2005;38(9):1798 806. [78] Niu J, Liu J, Zheng Y, Ran L, Chang Z. Are arch-conforming insoles a good fit for diabetic foot? Insole customized design by using finite element analysis. Hum Factors Ergon Manuf 2020;30(4):303 10. [79] Gu Y, Li J, Ren X, Lake MJ, Zeng Y. Heel skin stiffness effect on the hind foot biomechanics during heel strike. Skin Res Technol 2010;16(3): 291 6. [80] Sun JH, Cheng BK, Zheng YP, Huang YP, Leung JY, Cheing GL. Changes in the thickness and stiffness of plantar soft tissues in people with diabetic peripheral neuropathy. Arch Phys Med Rehabil 2011;92(9):1484 9. [81] Kattan KR. Thickening of the heel-pad associated with long-term Dilantin therapy. Am J Roentgenol Radium Ther Nucl Med 1975;124(1): 52 6. [82] Nack JD, Phillips RD. Shock absorption. Clin Podiatr Med Surg 1990;7(2):391 7. [83] Levy AS, Berkowitz R, Franklin P, Corbett M, Whitelaw GP. Magnetic resonance imaging evaluation of calcaneal fat pads in patients with os calcis fractures. Foot Ankle 1992;13(2):57 62.

Chapter 9

Multisegment Foot Models Amanda Stone1,2 1

RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 2ARCCA,

Incorporated, Seattle, WA, United States

Abstract Multi segment foot models (MFM) provide a means of dissecting the intricate kinematics and kinetics of the foot not able to be captured with single-segment foot models. Dozens of MFM have been published to date, some more successful and popular than others. This chapter discusses the important elements that one must consider when designing or selecting a published MFM, including the number of segments, bones being modeled, marker locations, and offset values. A review and comparison of the most popular models is also provided, namely the Milwaukee, Leardini/Rizzoli, Oxford, MacWilliams, and Bruening MFM. Finally, limitations and avenues for future research are highlighted.

9.1

Basic principles of multisegment foot models

9.1.1 Overview of motion capture Understanding how people move has been of interest to clinicians and researchers for decades. Studying the forces and motions that a person puts their body through often informs secondary factors, such as quality of life, risk of injury or re-injury, and life expectancy. The quantification of motion has undergone dramatic transformation over the last thirty years as motion capture technology continues to develop. Marker-based motion capture coupled with additional technology, such as force plates, have been around for quite some time, but modern software techniques and computing power have revolutionized the way kinematics and kinetics are measured. Reflective markers can be attached to the body in several ways, including directly on the skin, clustered on a skinmounted rigid plate, or directly attached to bones via a pin. Bone pins are the gold standard in terms of bone motion measurement [1,2], however require special training and introduce secondary complications such as infection and localized pain to the participant. As such, they will always remain a research tool and will never be used clinically. Skin-mounted markers are intended to represent the bone that lies underneath the skin, and are typically placed on bony landmarks for consistent and repeatable placement [3]. Despite careful placement, the skin slides over the underlying bone in dynamic situations (e.g., gait). This soft tissue artifact varies in severity across different bones and different areas of the body. Ideally, the range of motion of the joint of interest is substantially greater than the amount of soft tissue artifact to mitigate its impact. For instance, the knee joint’s range of motion while walking expands across 60 degrees; when this motion is translated to the markers mounted on the tibia, this equates to a minimum of 200 mm of motion. Comparatively, the soft tissue artifact of the tibia markers can be as large as 20 mm, or approximately 10% of the marker’s overall motion [4,5]. To quantify a person’s biomechanics, the body is divided into segments and the relationships between segments are quantified. For example, the relationship between the pelvis and the femur can be quantified as the hip angle. Segments are defined based on the researcher’s goals, and can be defined as either a single bone (e.g., the thigh segment is defined as the femur) or a combination of bones (e.g., the lower leg is defined as the tibia and fibula). Threedimensional body motion is captured via a minimum of three markers on each of the segments utilized to calculate the Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00018-4 © 2023 Elsevier Inc. All rights reserved.

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joint angle of interest. For example, markers placed on the pelvis, femur, and lower leg can be combined to calculate hip and knee motion during gait.

9.1.2 Why use a multisegment foot model? All models must straddle the gap between variables of interest, necessary assumptions, and available resources. Perhaps a researcher is only interested in frontal knee moments and thus does not require full body modeling, or they do not have enough markers to apply a full body model. As will be discussed later, each researcher must determine their own personal interests and where they feel comfortable drawing the line between the advantages and disadvantages of different models. Single-segment foot models are used by researchers and clinicians alike for ease of application and interpretation. While this may be suitable for some who are most concerned with the biomechanics of the hip or knee, single-segment foot models do not provide the level of detail needed to thoroughly represent the complex biomechanics of the foot and ankle. Single-segment foot models are a dramatic simplification as the human foot typically contains 28 bones. Additionally, some research questions, such as the kinematic differences between persons with pes cavus versus pes planus, cannot be answered with a single-segment foot model. Multisegment foot models (MFMs) provide a means of decomposing the foot’s intricacy and exploring the biomechanics within the foot. Studies comparing the kinematic and kinetic outcomes derived from single- versus MFM have illustrated drastic differences depending on the modeling technique [6 8]. The role of the ankle is often overestimated in single-segment models [6,7,9,10], with sagittal ankle power overestimated by as much as 53% [9,10]. This is partially because all measured motion is attributed to the ankle joint in a single-segment model, and joints within the foot which contribute to that measured motion are masked [9]. For example, the midfoot undergoes a greater range of motion than the ankle joint during gait [11,12]. Furthermore, direction of motion can change based on the model used; when analyzing gait, Pothrat et al. calculated persons with pes planus exhibited ankle plantarflexion at heel strike when using a singlesegment foot model, and ankle dorsiflexion when using a MFM [8]. These direct comparisons between single- and MFMs illustrate the value and necessity of modeling these distal segments.

9.2

Selecting an appropriate multisegment foot model

As with any modeling technique, there is a balance between what information is measured or calculated and what information is assumed or ignored. While theoretically possible, measuring every joint in the human foot would lead to lengthy data collection times, extensive data processing, and an overwhelming amount of data to interpret. Each researcher must ask themselves what question they are investigating and what outcomes they can accurately measure. The following subsections discuss the most important considerations when selecting a MFM. While they are presented separately, a number of these considerations are interdependent. For instance, the degrees of freedom of a model is dependent on the number of palpable landmarks and the amount of surface area on the foot [13]. Likewise, the number, quality, and location of available cameras dictates the number of markers and their proximity to each other [13,14].

9.2.1 Segments and bone groupings The number of segments and which bones are included in those segments is arguably the most pivotal decision when selecting an existing MFMs or creating a new one. While the scientific community generally agrees the foot should be broken into segments [15,16], there is discord with regard to what bones can and should be modeled, and how they should be grouped together. Although MFMs mitigate many of the limitations single-segment models introduce, it is inevitable that certain joints’ true motion will be masked. Some MFMs, such as the Oxford MFM, break the foot into a hindfoot and a forefoot, with the midfoot lacking representation [17,18]; other MFMs, including the Rizzoli MFM, group the five bones that compose the midfoot into one segment [14,19]. Any bones grouped together are assumed to move as one rigid segment, so segment and bone grouping selection must consider this limitation. Nester et al. point out the decision to model neighboring bones as one rigid segment is nontrivial and should not be taken lightly [20]. Bone pin studies have demonstrated relative motion between adjacent joints in a number of interhindfoot and -midfoot joints, further emphasizing that all joints contribute to motion of the foot [11,20]. As bone groupings are unavoidable, care must be taken when deciding which bones to group together. Nester et al. tested different combinations of bones to see which best represented true foot motion, and proposed specific bones be grouped together, such as the navicular and cuboid or the cuneiforms and metatarsals [13]. The goal is to model what is most important for

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your particular study. For example, modeling a hallux segment may be instrumental in assessing foot biomechanics in persons with hallux valgus; or the midfoot could provide newfound insights about the differences between symptomatic and asymptomatic pes planus.

9.2.2 Marker type and placement Reflective markers are placed on bony landmarks to create a direct link between the skin-mounted marker and the underlying anatomy that the marker is intended to represent. By placing markers on bony easily palpable landmarks—such as the navicular apex—marker placement, and thus foot models, can be more consistent between examiners, participants, and sessions. One of the major differences between full body modeling and foot modeling, however, is the size of the bones and joints being modeled. The tibia, for example, is almost 8 times longer than the first metatarsal (400 mm vs 50 mm on average) [21 24]. Smaller bones equate to a smaller number of palpable bony landmarks and a reduction in the available skin surface area. The majority of foot bones do not physically have enough skin surface area to facilitate the minimum number of markers needed to model the joint in three dimensions. Moreover, markers that are closer together are both more difficult to track and more sensitive to errors in placement. Researchers combat these obstacles in several ways, including modeling fewer segments or utilizing cluster wands that have a smaller footprint.

9.2.3 Coordinate systems Anatomically-based coordinate systems are the primary method utilized in MFMs, although several models employ technically-based coordinate systems. In an anatomical coordinate system, markers are placed on bony landmarks and are directly tracked throughout the static and dynamic trials [9,10,14,17,18,25,26]; the rotations between coordinate systems have anatomical meaning. A technical coordinate system typically utilizes markers that are arbitrarily placed (e.g., cluster plates); anatomical descriptions are only possible if the arbitrary systems are linked to bony landmarks during a static trial. The original Leardini MFM applies this technique using a pointer calibration to identify anatomical landmarks [19], and only the cluster plates are tracked during dynamic trials. A minimum of three noncollinear markers per segment are required for three-dimensional motion analysis. From these markers, a local coordinate system is generated and three-dimensional joint rotations can be calculated [27]. Whatever the reason, whether due to available anatomical landmarks or the number of segments being modeled, threedimensional motion may not be possible for all segments in a given MFM. In these circumstances, planar angles, or two-dimensional angles, are utilized, as is evident amongst a number of MFMs [14,28,29]. Two markers are connected to create a line segment and the relationship between line segments is quantified to represent two-dimensional motion.

9.2.4 Offsets The difference between the intermarker distance and the range of motion that the joints of interest go through is smaller in MFMs than in full body models. Because of this, MFMs are more sensitive to marker placement accuracy, where illplaced markers can generate large errors and confound the results. Offset values have been implemented to normalize data to a standardized pose, such as normal standing [14] or subtalar neutral [30]. Subtracting a known value from dynamic trials alleviates the effect of marker placement accuracy and produces lower variability when compared to absolute angles (i.e., nonoffset data) [31]. Unfortunately, applying an offset value can mask underlying anatomy, especially in persons with foot deformities or abnormalities, where the differences in measured motion is due to true anatomical differences and not errors in marker placement [32]. In certain populations, such as healthy individuals, this may not be as great of a concern, again further emphasizing the importance of selecting a MFM based on the researcher’s unique situation. Regardless of whether an offset value is used, kinematics will be described with a certain value specified as “neutral,” or 0 degrees, and subsequent motion will be based off that value. For example, sagittal motion is most often described with 0 degrees representing the ankle at 90 degrees, with motion greater than 90 degrees described as plantarflexion and motion less than 90 degrees described as dorsiflexion. However, some populations, such as those with severe equinus contracture, may not be able to achieve 90 degrees at the ankle. Researchers must consider how to account for these factors when designing and/or implementing a MFM.

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9.2.5 Standardized description of multisegment foot models Review articles summarizing the available MFMs have identified over 30 MFMs in the published literature [33 37], which is likely a small representation of the true number when considering the vast amount of in-house MFMs that do not gain widespread traction. With no standard for presenting MFM protocols and methodologies, the information provided from model to model is divergent. This causes issues in reproducibility of results or when trying to compare MFM kinematics and kinetics amongst studies. Different ventures have been attempted by the scientific community to provide direction and instruction for a uniform approach. Bishop et al. provided a detailed list of five standards each MFM publication should include or provide reference to; specifically, each article should report: 1. 2. 3. 4. 5.

Location, accuracy, and reliability of marker placement Definition of segment(s) Definition of segment coordinate systems Definition of joint parameters Reliability of joint kinematics

9.3

Review of current multisegment foot models

This section reviews the most popular MFMs used, namely the Milwaukee, Leardini/Rizzoli, Oxford, MacWilliams, and Bruening models. The modeled segments, grouped bones, and calculated biomechanics will be discussed for each model. There are also a plethora of review articles which provide a more detailed and thorough assessment of published MFMs [33 37]. Terminology varies across studies (e.g., hindfoot vs rearfoot, eversion vs pronation); the terminology has been modified from its original text where possible to facilitate direct comparison across models, with the original terminology listed in parentheticals.

9.3.1 Milwaukee kinematic model The Milwaukee MFM is a four-segment foot model that comprises 12 markers which define the (1) tibia and fibula, (2) calcaneus, talus, and navicular, (3) cuneiforms, cuboid, and metatarsals, and (4) proximal phalanx of the hallux [25] (Fig. 9.1). Three markers were placed within each segment, although not every bone in the segment was identified. For example, the calcaneus/talus/navicular segment was modeled via three markers on the calcaneus. Furthermore, the cuneiforms/cuboid/metatarsals segment was modeled using one marker on the first metatarsal and two markers on the fifth metatarsal. The hallux segment was modeled using a rigid marker triad which extended from the skin surface. Radiographs were employed to better align marker-based coordinates with the underlying bone anatomy. Threedimensional kinematics were calculated for leg-to-lab (originally tibia-to-lab), hindfoot-to-leg (originally hindfoot-totibia), forefoot-to-hindfoot, and hallux-to-forefoot joint rotations. Long et al. assessed the model’s repeatability and found low measurement error within and across sites and sessions with the largest source of variability across subjects [38]. From a clinical perspective, the Milwaukee MFM has been applied to a wide range of ages and conditions including, but not limited to children with equinovarus [39], adults with posterior tibial tendon dysfunction [40 42], adults with hallux rigidus or hallux valgus [43 46], adults with clubfoot [47], adults with ankle osteoarthritis [48,49], and adults with rheumatoid arthritis [50].

FIGURE 9.1 Diagram of the foot with reflective marker placement for the Milwaukee MFM.

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9.3.2 Leardini/Rizzoli kinematic model The Leardini MFM was first described as a five-segment foot model consisting of the (1) tibia and fibula, (2) calcaneus, (3) navicular, lateral, middle, and medial cuneiforms, and cuboid, (4) first metatarsal, and (5) proximal phalanx of the hallux [19] (Fig. 9.2). No known models prior to this had quantified midfoot motion independent of the hindfoot or forefoot. The model utilized five rigid cluster plates, each with four reflective markers, mounted on the five segments. A pointer calibration was used to identify the location of 16 anatomical landmarks and calculate three-dimensional joint rotations for the hindfoot-to-leg (originally calcaneus-to-tibia), midfoot-to-hindfoot (originally midfoot-to-calcaneus), forefoot-to-midfoot (originally metatarsals-to-midfoot), and hallux-to-forefoot (originally hallux-to-metatarsals). The static standing foot posture was also quantified and subtracted from the calculated joint rotations in all dynamic trials. Numerous studies assessed clinical populations with this original model, including high- and low-arched athletes [51] and persons with talocalcaneal coalition [52]. The model was modified and later renamed the Rizzoli MFM, still containing five segments but with the forefoot (originally metatarsal) segment being extended to include the first through fifth metatarsals, the hallux segment removed, and an overall foot segment added [14] (Fig. 9.3). The Rizzoli model also discontinued the use of cluster plates and pointer calibration. This model utilized 15 markers placed directly on the skin to calculate three-dimensional joint rotations: foot-to-leg (originally foot-to-shank), hindfoot-to-leg (originally calcaneus-to-shank), midfoot-to-hindfoot (originally midfoot-to-calcaneus), forefoot-to-midfoot (originally metatarsal-to-midfoot), and forefoot-to-hindfoot (originally metatarsal-to-calcaneus). Two-dimensional planar angles were also calculated, divided into five sagittal and three transverse plane

FIGURE 9.2 Diagram of the foot with cluster plate placement (top) and pointer calibration locations (bottom) for the original Leardini MFM.

FIGURE 9.3 Diagram of the foot with reflective marker placement for the Rizzoli MFM.

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angles. Static standing foot posture was only subtracted from the joint rotations, not the planar angles. Studies assessing the variability and repeatability of the Rizzoli MFM found favorable findings ranging from moderate to excellent [28,31,32,53 56]. Numerous studies emphasized the importance of careful training and expertise with respect to repeatability and variability [31,53,54]. The Rizzoli MFM is the most repeatable and reproducible compared to other models, one of which being the Oxford MFM [32]. Portinaro et al. introduced a modified version of the Rizzoli MFM which focused on increasing consistency with clinical findings [28]. The modifications entailed an additional marker attached to the posterior inferior calcaneus to enhance frontal plane orientation, redefinition of the medial longitudinal arch calculation, and reintroduction of a hallux segment using a novel three-marker definition. The Rizzoli MFM has been applied in children with flatfeet [28,57], healthy adults [58 61], and adults with a wide variety of conditions including plantar fasciitis [62], diabetes and neuropathy [63], chronic ankle instability [64,65], tibialis posterior tendon dysfunction [66], knee osteoarthritis [67], and forefoot varus [68].

9.3.3 Oxford kinematic model The Oxford MFM is a four-segment foot model, which includes 18 markers to define the (1) tibia and fibula, (2) calcaneus and talus, (3) first through fifth metatarsals, and (4) proximal phalanx of hallux [17] (Fig. 9.4). Sagittal, coronal, and transverse plane kinematics were calculated for the leg-to-lab (originally tibia-to-floor), hindfoot-to-leg (originally hindfoot-to-tibia), forefoot-to-hindfoot, and hallux-to-forefoot. Clinical studies utilized this initial Oxford MFM in a breadth of individuals, such as children with equinovarus [69] or flatfeet [70], and adults with calcaneal fracture [71], subtalar arthrodesis [71], hallux rigidus [72], rheumatoid arthritis [16,73,74], and high- and low-arched athletes [51]. Stebbins et al. released modifications to the Oxford MFM to better accommodate children and persons with foot deformities [18] (Fig. 9.5). The updated model redefined the leg (originally tibia) coordinate system by adding a lateral epicondyle marker, altered the hindfoot calculations to be independent of adjacent segments, moved the proximal first metatarsal marker slightly dorsal, and replaced the hallux triad with a single marker, resulting in 17 markers in the modified model. Three-dimensional joint rotations were calculated for the leg-to-upper leg (originally tibia-to-femur), hindfoot-to-leg (originally hindfoot-to-tibia), forefoot-to-hindfoot, and forefoot-to-leg (originally forefoot-to-tibia). Sagittal hallux-to-forefoot kinematics and dynamic arch height were also quantified. This updated model was subject to several repeatability studies, which generally found high repeatability with regard to sessions and examiners [32,55,75 81].

FIGURE 9.4 Diagram of the foot with reflective marker placement for the Oxford MFM.

FIGURE 9.5 Diagram of the foot with reflective marker placement for the modified Oxford MFM.

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FIGURE 9.6 Diagram of the foot with reflective marker placement for the MacWilliams/Kinfoot MFM.

While the modified Oxford MFM is sensitive to marker misplacement [77], it is not as sensitive as other MFMs when directly compared [32]. Stebbins et al. [18] made no reference to static position subtraction, however Wright et al. and Milner and Brindle demonstrated the stability of relative values compared to absolute values, with a reference to neutral stance reducing error and variability [76,80]. The modified Oxford MFM has been used extensively to assess foot kinematics in healthy adults [82] and persons with flatfeet [83 85], forefoot varus [86], ankle fracture [87], hallux valgus [88], cerebral palsy [81,89], clubfoot [81,90], excessively-everted feet [91], patellofemoral pain syndrome [92 94], pes planovalgus [95], and midfoot pain [96].

9.3.4 MacWilliams/Kinfoot kinetic model By the early 2000s, dozens of kinematic MFMs had been described in the scientific literature; however, MacWilliams et al. were the first to develop a model that also quantified three-dimensional foot kinetics [9]. The foot model, informally termed Kinfoot, comprises 19 markers that represent nine segments: (1) tibia and fibula, (2) talus, navicular, and cuneiforms, (3) cuboid, (4) calcaneus, (5) lateral forefoot, (6) medial forefoot, (7) lateral toes, (8) medial toes, and (9) hallux (Fig. 9.6). Kinetic data were acquired via utilization of a pressure mat and force plate during two separate trials. Measured forces were then distributed across six subdivided sections (hindfoot, medial forefoot, lateral forefoot, first toe, second/third toes, and fourth/fifth toes). Sagittal, coronal, and transverse angles were quantified for three hindfoot joints (talocrural, subtalar, and calcaneocuboid), two midfoot joints (medial and lateral tarsometatarsal), and three toe joints (hallux metatarsal, and medial and lateral metatarsal phalangeal). Three-dimensional moments and powers were also calculated for each of the previously listed joints. The subtalar and calcaneocuboid joints were the least repeatable regarding kinematic parameters, while the lateral tarsometatarsal and lateral toes were the least repeatable with respect to kinetic parameters [9]. Kinematically, the MacWilliams MFM was moderately repeatable, but exceeded a mean error of 5 degrees over repeated sessions [55].

9.3.5 Bruening kinetic model Bruening et al. outlined several challenges regarding kinetic MFMs in general, in addition to previously developed models, and sought to create a foot model that mitigated these limitations and shortcomings [10,26]. Their model consisted of 19 markers which defined five segments: (1) leg (originally shank), (2) hindfoot, (3) forefoot, (4) hallux, and (5) overall foot (Fig. 9.7). Hindfoot-to-leg (originally hindfoot-to-shank, that is, ankle complex, a combination of the talocrural and subtalar joints), forefoot-to-hindfoot (i.e., midtarsal joint), and hallux-to-forefoot (i.e., metatarsophalangeal joint) kinematics and kinetics were calculated for all three planes of motion aside from the hallux-to-forefoot kinematics which did not include the coronal plane. Ground reaction forces were captured with the foot spanned across two force plates in staggered positions to isolate the midtarsal or metatarsophalangeal joint separately, and ultimately calculate joint moments and powers. This model has been utilized to better understand intersegmental power within the foot in healthy adults [97,98].

9.3.6 Direct comparison of current multisegment foot models A comparison of the Milwaukee, Leardini/Rizzoli, Oxford, MacWilliams, and Bruening MFMs illustrates the similarities and differences between them (Table 9.1). The Milwaukee, Rizzoli, Oxford, MacWilliams, and Bruening MFMs provide clear descriptions of their marker locations. The original Leardini model provides a clear description of the pointer locations, but does not dictate where the cluster plates should be mounted.

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FIGURE 9.7 Diagram of the foot with reflective marker placement for the Bruening MFM.

TABLE 9.1 Description of some of the more common multisegment foot models. MFM

#Markers

Segments

Milwaukee [25]

12

1. 2. 3. 4.

Leardini [19]

20

1. Leg (tibia, fibula)a 2. Hindfoot (calcaneus)a 3. Midfoot (navicular, lateral/middle/medial cuneiforms, cuboid) 4. Forefoot (1st metatarsal)a 5. Hallux (proximal phalanx)

Hindfoot-to-lega Midfoot-to-hindfoota Forefoot-to-midfoota Hallux-to-forefoot

Rizzoli [14]

15

1. Leg (tibia, fibula)a 2. Hindfoot (calcaneus)a 3. Midfoot (navicular, lateral/middle/medial cuneiforms, cuboid) 4. Forefoot (1st/2nd/3rd/4th/5th metatarsals)a 5. Foot

Foot-to-lega Hindfoot-to-lega Midfoot-to-hindfoota Forefoot-to-midfoota Forefoot-to-hindfoot

Original Oxford [17]

18

1. 2. 3. 4.

Modified Oxford [18]

17

Same as Original Oxford [17]

Leg-to-upper lega Hindfoot-to-lega Forefoot-to-hindfoota Forefoot-to-leg

MacWilliams [9]

19

1. 2. 3. 4. 5. 6. 7. 8. 9.

Talocrural Subtalar Calcaneocuboid Medial tarsometatarsal Lateral tarsometatarsal Hallux metatarsal Medial metatarsophalangeal Lateral metatarsophalangeal

Bruening [10,26]

19

1. Leg (tibia/fibula)a 2. Hindfoot (calcaneus, talus) 3. Forefoot (navicular, cuboid, cuneiforms, metatarsals) 4. Hallux (proximal/distal phalanges) 5. Foot

a

Leg (tibia, fibula)a Hindfoot (calcaneus, talus, navicular) Forefoot (cuneiforms, cuboid, metatarsals) Hallux (proximal phalanx)

Leg (tibia, fibula)a Hindfoot (calcaneus, talus) Forefoot (1st/2nd/3rd/4th/5th metatarsals) Hallux (proximal phalanx)

Leg (tibia, fibula) Hindfoot (calcaneus) Midfoot (talus, navicular, cuneiform) Cuboid Medial forefoot Lateral forefoot Medial toes (2nd/3rd metatarsals) Lateral toes (4th/5th metatarsals) Hallux

Terminology was changed from the original publication to facilitate direct comparison across models.

Joint outcomes Leg-to-laba Hindfoot-to-lega Forefoot-to-hindfoot Hallux-to-forefoot

Leg-to-laba Hindfoot-to-lega Forefoot-to-hindfoot Hallux-to-forefoot

Hindfoot-to-lega Forefoot-to-hindfoot Hallux-to-forefoot

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The modified Oxford MFM is the only model to include an upper leg segment (i.e., femur) in the model and calculate a leg-to-femur three-dimensional angle. The leg segment is uniform across all MFMs, including both the tibia and fibula, although the number and placement of markers vary. Three-dimensional leg-to-lab angles are calculated in the Milwaukee and original Oxford models. While all MFMs include a hindfoot segment, different bones are grouped depending on the model. The Milwaukee, Oxford, and Bruening MFMs only place markers on the calcaneus, but include other bones, such as the talus or navicular in the hindfoot segment; the original Leardini, Rizzoli, and MacWilliams MFMs limit the hindfoot segment to only include the calcaneus. All MFMs also quantify hindfoot-to-leg biomechanics with the exception of the MacWilliams MFM. The MacWilliams model quantifies three-dimensional talus-to-leg and calcaneus-to-talus kinematics and kinetics despite not having specified a talus segment (the talus is grouped with the navicular and cuneiforms to compose the midfoot segment). The Rizzoli and Bruening MFMs also modeled the foot as a single-segment for comparison purposes and provided three-dimensional angles, moments, and powers for the foot-to-leg. The midfoot is modeled in the original Leardini and Rizzoli MFMs; it includes the navicular, cuneiforms, and cuboid. In the original Leardini model, the cluster plate generally overlaps these bones, however in the Rizzoli model, only a single marker is placed on the navicular apex. Both models calculate midfoot-to-hindfoot kinematics. In addition to threedimensional joint rotations, the Rizzoli MFM also proposes a number of two-dimensional, planar angles to describe midfoot motion. The MacWilliams MFM groups the talus, navicular, and cuneiforms together to represent the midfoot, however no markers are placed on any of the bones in this region. The MacWilliams MFM also models a cuboid segment and calculates three-dimensional cuboid-to-calcaneus kinematics and kinetics, but lacks a designated marker on the cuboid. Moving on to the forefoot, all models include a forefoot segment, but similar to the hindfoot, vary in terms of which bones compose the segment. The Milwaukee MFM forefoot segment involves the cuneiforms, cuboid, and metatarsals, but only places markers on the metatarsals. The original Leardini MFM only models the first metatarsal, however it is difficult to discern if the cluster plate is small enough to isolate this single metatarsal. The Rizzoli and Oxford MFMs include all five metatarsals in the forefoot segment which is modeled by markers placed on the first, second, and fifth metatarsals. Forefoot-to-hindfoot kinematics are provided by the Milwaukee, Rizzoli, Oxford, and Bruening MFMs. Three-dimensional moments and powers are also modeled in the Bruening MFM for the forefoot-to-hindfoot. Additionally, forefoot-to-midfoot joint rotations are outlined by the original Leardini and Rizzoli models. The Rizzoli MFM also calculates a few two-dimensional angles to describe motion between individual metatarsals. The forefoot is broken into medial and lateral components in the MacWilliams MFM, however there is no clear definition of what bones these segments represent. Furthermore, three-dimensional medial and lateral tarsometatarsal kinematics and kinetics are presented with no clear description of the derived segments. The forefoot segment of the Bruening MFM includes the navicular, cuboid, cuneiform, and all metatarsals and includes markers on the navicular, cuboid, and first, second, and fifth metatarsals. Single markers, cluster plates, and marker wands are all used to model the hallux depending on the MFM. The original Leardini model uses a four-marker cluster plate, although, similar to the first metatarsal, it is unclear if the plate is small enough to isolate hallux motion without incorporating the first metatarsal or distal phalanx of the hallux. While the Rizzoli MFM removes the hallux segment from the model, Portinaro et al.’s proposed modification reintroduces the hallux segment, modeled as a single marker [28]. The Bruening MFM also implements a single marker on the hallux. The Milwaukee, Oxford, and MacWilliams MFMs utilize a three-marker wand mounted to the hallux, although this is changed to a single marker in the modified Oxford MFM. Hallux-to-forefoot three-dimensional joint rotations are quantified in the Milwaukee, original Leardini, Portinaro et al., modified Rizzoli, original Oxford, and MacWilliams MFMs. The MacWilliams model also provides three-dimensional hallux-to-forefoot moments and powers. The Stebbins et al.-modified Oxford model included sagittal hallux-to-forefoot angles, and the Bruening model included sagittal and transverse halluxto-foot angles. The Bruening model also included three-dimensional moments and powers for the hallux-to-foot. Similar to the forefoot, the toes are broken into medial and lateral segments by the MacWilliams MFM, though no description of the representative bones is provided. Three-dimensional medial and lateral metatarsophalangeal kinematics and kinetics are presented with no clear description of which segments the joint rotations are composed of.

9.4

Applications, considerations, and limitations

9.4.1 Clinical applications MFMs provide the means to assess intricate foot motion which is masked using single-segment foot models. Clinical populations in particular may exhibit drastic changes within the joints of the foot. A breadth of maladies have been

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studied thus far, including equinovarus [39], chronic ankle instability [6,64,65], ankle osteoarthritis [48,49,99 103], knee osteoarthritis [67], clubfoot [47,81,90], foot type (e.g., pes planus or cavus) [28,57,83 85,104,105], hallux rigidus or hallux valgus [43 46,88], rheumatoid arthritis [50], talocalcaneal coalition [52], diabetes [63,106,107], posterior tibial tendon dysfunction [40 42,66], patellofemoral pain syndrome [92 94], plantar fasciitis [62], and cerebral palsy [81,89,108].

9.4.2 Sources of error No model is without error. The most common sources of error for MFMs are soft tissue artifact (i.e., skin motion artifact), limited skin surface area, and violation of rigid body assumptions. While these sources are known, it is difficult to quantify their relative contributions to the overall error as many are interdependent. Furthermore, while it is important to consider these sources of error when selecting a MFM, each study will have different requirements and a given MFM and its inherent error may be satisfactory for one study but inadequate for another; the amount of acceptable error will depend on the aims of the study, the population being investigated, and the bones and joints of interest [13]. As discussed previously, soft tissue artifact occurs when the skin, muscle, and other soft tissues move independent of the bones they are enveloping. As skin-mounted markers and clusters are intended to mirror the motion of the foot bones that lie under the skin, any deviation would result in an inaccurate representation of foot motion. When assessing gait, even contact with the ground during heel strike can introduce soft tissue movement due to the impact force. Numerous studies have quantified soft tissue artifact and identified the estimated error that arises at each modeled segment and joint [2,20,32,109 114]. The malleoli and calcaneus markers had the largest artifact in the Rizzoli [14] and Oxford MFMs [18] when compared to CT scans of the foot, which affected the hindfoot-to-leg joint angles [114]. Radiographic images illustrated artifact between marker wands and underlying bone could be as large as 16 mm (navicular) [113]. When examined as a whole, the soft tissue artifact observed in the foot lacks consistency across segments; no single plane of motion or joint consistently generated the largest error [20], making it difficult to implement a blanketed solution. While large magnitudes of soft tissue artifact contribute a significant amount of error, this is compounded when coupled with the small ranges of motion observed in foot joints [32]. Inherent anatomy of the foot facilitates erroneous measurements due to limited bony landmarks and narrow spacing between markers. Caravaggi et al. noted greater variability was expected in MFMs compared to full body models due to smaller bony landmarks and less marker separation [53]. The midfoot, for example, often exhibits the highest variability compared to other segments, which may be due to the proximity of markers [32]. Close proximity of markers, along with slight deviations in marker placement, can result in large divergences in subsequently calculated anatomical coordinate systems and joint rotations [18,53]. Landmark availability is also limited on the foot. Some bony landmarks may be moved or absent due to foot deformities [59,115,116]. Furthermore, certain bones, such as the talus, are difficult to track at the skin level without invasive methods such as bone pins; this leads to the talus being grouped with neighboring bones, such as the calcaneus, despite motion occuring between these two bones. One of the unavoidable drawbacks of biomechanical modeling is the inevitability of violating the rigid body assumption. When dividing the foot into smaller segments, the assumption is made that no movement occurs between individual foot bones within a segment; numerous studies have demonstrated motion between all measured bones in the foot, highlighting a direct conflict with the rigid body assumption [11,110]. For instance, the cuboid and navicular are grouped together as in a single segment in the original Leardini, Rizzoli, and Bruening MFMs despite known motion between these bones [11]. This discrepancy then leads to measured motion being attributed to different joints than where it is truly occurring. While it can be difficult to pinpoint how much relative error different sources (e.g., soft tissue artifact) contribute, hindfoot-to-leg joint rotations most closely represent true bone motion compared to other joints [20,112]. Thus, it is likely that violation of the rigid body assumption is the greatest contributor to measured error [20]. Kinetic MFMs must face newfound challenges and obstacles that are not present in kinematics models. MacWilliams et al. outline a kinetic model, which utilizes a pressure mat and force plate to obtain forces that are then distributed across foot segments to generate joint three-dimensional moments and powers [9]. There is no clear agreement within the scientific literature, however, with how to interpret and apply shear forces in a MFM. Yavuz et al. illustrated peak shear force and peak vertical force can occur at different locations on the foot, and that it is very difficult to predict shear force based on vertical pressure [117]. Furthermore, isolating and measuring the forces acting upon a given segment is challenging given the narrow margin of space between segments. Bruening et al. describe a simplistic method of measuring forces by staggering foot placement across two force plates. Even then, application of force to the

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modeled foot segments requires additional assumptions to be made. Kinetic MFMs are still controversial due to violation of the rigid body assumption and its effect on anatomical coordinate systems, axis of rotations, and subsequent kinetic calculations [20].

9.5

Areas of future biomechanical research

While MFM modeling has improved since this technique was introduced decades ago, there are still many avenues and possibilities for future development and research. While many studies have assessed the repeatability of a given MFM, studies evaluating the validity and accuracy of a model are lacking. Radiographs and CT scans have been utilized previously in the extant literature to directly compare bone positioning to marker placement, though these comparisons are often made in static or quasistatic poses. Biplane fluoroscopy would provide a means of measuring foot bone motion in a dynamic setting and increase the functional relevance of the study’s findings. Moving beyond reflective markers and infrared cameras, inertial measurement unit (IMU)-based systems could take MFMs into the real-world and escape the constraints of a lab. Depending on the size of the IMU, more segments could be modeled as marker spacing with respect to camera angles would no longer be a consideration. Another channel to consider is the development or assessment of MFMs with respect to footwear. Shoes are essential in certain studies, such as those measuring the effect of orthoses, but introduce new barriers when it comes to palpating bony landmarks and indirectly quantifying bone motion. Some studies have placed markers on top of the shoe [118], cut out holes in the shoe to be able to directly place markers on the skin [61,119,120], or used sandals and placed markers around the straps [104]. It is unclear, however, how these approaches affect the accuracy of the data, or in the case of cutting holes in the shoe, the function of footwear. Lastly, there is still much to be determined when it comes to kinetic MFMs. Kinematic MFMs have increased in terms of popularity and number of available models, but the kinetic equivalent has not. New measurement techniques (e.g., shear sensors) and the impact of specific assumptions can be more widely investigated, potentially leading to more robust kinetic models.

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Foot motion in children shoes—a comparison of barefoot walking with shod walking in conventional and flexible shoes. Gait Posture 2008;27:51 9. Available from: https://doi.org/10.1016/j.gaitpost.2007.01.005. [120] Shultz R, Birmingham TB, Jenkyn TR. Differences in neutral foot positions when measured barefoot compared to in shoes with varying stiffnesses. Med Eng Phys 2011;33:1309 13. Available from: https://doi.org/10.1016/j.medengphy.2011.05.009.

Chapter 10

Invasive Techniques for Studying Foot and Ankle Kinematics* Arne Lundberg1 and Anton Arndt1,2 1

Department of Clinical Science, Intervention and Technology, Karolinska Institute, Stockholm, Sweden, 2The Swedish School of Sport and Health

Sciences (GIH), Stockholm, Sweden

Abstract Biomechanists have only relatively and recently realized the importance of representing the foot in terms of multiple segments instead of a single rigid segment. The necessity arose from prosthetic advances and from a demand for more attenuated information for describing sport biomechanics. However, a kinematic differentiation between the many small, irregularly shaped bone segments within the foot poses considerable methodological challenges. Conventional skin marker-based optoelectronic motion analysis is extremely limited as the skin markers do not accurately represent individual skeletal segments. Numerous studies have been performed in which invasive techniques were applied to overcome this problem. Radiostereometric analysis (RSA) involves the insertion of radio opaque metal beads in the segments of interest and the subsequent three-dimensional analysis of the segments using biplanar X-ray. Intracortical pins can be directly inserted into the relevant bones, and standard reflective markers are attached to the protruding pin, permitting dynamics kinematic analysis with standard optoelectronic systems. Such studies have provided seminal basic research data describing the motion of the foot and ankle. From radiostereometric studies, we now have an accurate description of the ankle joint rotation axes at different positions, which is an important information for ankle prosthetic surgery and design. Similarly, both RSA and intracortical pin studies have provided information regarding the movement of the talus during gait, which is not possible using skin marker-based methods. Differences in range of motion in joints of the foot between walking and running have been described. In more applied studies, intracortical pins have been applied to investigate the effects on intrinsic foot kinematics when using foot orthoses inside shoes. The results of this body of research are important for many clinicians and scientists who may not have the possibility of conducting invasive experiments, but can apply the available data in validation of more simplified foot models.

10.1

Introduction

This chapter provides a historical perspective of the identified necessity of and the development of invasive methods for describing ankle and intrinsic foot kinematics. Biomechanists were for a surprisingly long time satisfied with a simplified model of the foot, often as a single rigid segment, which was not sufficiently complex for explaining many scientific questions. The major studies using the invasive techniques radiostereometric analysis (RSA) and intracortical pins to contribute to this field of knowledge are outlined. The advantages of these techniques over skin marker based methods are described but a section is also included summarizing the limitations and exploring possible future directions in foot and ankle kinematics methodology. Finally, an appendix is provided in which the authors have attempted to share their experience of the invasive methods and indicate critical points of which one must be aware when inserting either X-ray markers or intracortical pins in the complex skeletal segments of the foot and ankle. *This chapter is dedicated to the memory of Alex Stacoff, who was one of the most driving and enthusiastic members of our group, and whose untimely death has meant that our knowledge of foot and ankle biomechanics is less than it might have been. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00017-2 © 2023 Elsevier Inc. All rights reserved.

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Early invasive studies of foot and ankle biomechanics

When scientific study of human movement was first performed in the late 19th and early 20th centuries, limited attention was usually spent on the foot and ankle. However, in the later part of the 20th century, many researchers started to consider the foot as a rigid block attached to the lower leg with a hinge. However, a need for a more detailed analysis was growing during that period, fueled by several specific research needs such as the following: (1) The need for improved prosthetics, which after World War II led to a surge of interest in lower limb biomechanics, largely within the US Department of Veterans Affairs. (2) The perceived need for better understanding of sports physiology. (3) The need to understand if the failure of early attempts at ankle joint replacement had partly biomechanical causes. Traditionally, biomechanical research (except for photography or film-based in vivo studies such as those by Muybridge or occasional invasive studies using bone pins, such as Levens et al. [1] and Close et al. [2]), was based on the examination of cadaveric specimens. While these studies have been very useful in many fields, they have also relied on the assumption that some obvious limitations do not influence the results in a major way. The most obvious limitations are: (1) The specimens examined are usually anatomically limited, meaning that the absence of structures outside of the specimen under investigation might have influenced the results (i.e., the specimen is rarely the whole-body or even whole-limb). (2) The input forces will have to be controlled as they are not the result of any natural activity. For vertical load, this may not be a major factor, but in joint kinematics (other than as induced by vertical load) the risk of introducing movement that would not occur normally may be considerable. (3) The procedure of preparing the specimens may affect their biomechanical properties. From the 1940s and 40 years onward, there was a large increase in the amount of information gathered mainly from the studies of cadaveric specimens. Some of these originated from the post-World War II upsurge in prosthetic research [3], some from more traditional academic sources [4]. Depending on the method used, results seemed to vary, particularly for talocrural joint kinematics. Some aspects, such as the transverse plane rotation in the lower limb during walking, were occasionally studied by invasive means [1] with results indicating that the transverse plane rotation is being transferred and transformed through the joints of the lower extremity, particularly those of the ankle and foot. To bridge the gap between in vitro studies and in vivo kinematics, several methods have been tried for reducing the errors that may be introduced by attempting to study bone kinematics using body surface information; this presentation will deal with some invasive applications that were specifically aimed at the foot and ankle in the absence of noninvasive alternatives. In the absence of reliable and accurate methods for noninvasive measurements in foot and ankle kinematics, it has been—and sometimes still is—necessary to insert or attach markers directly into bone to collect data representing true skeletal movement. One way to achieve this is RSA, which involves insertion of radio opaque markers (usually made of tantalum, which is a nonreactive and very dense metal, recently popular in implant manufacturing) into the bones, wait for any possible discomfort from the insertion to subside and then make dual radiographic exposures in a calibrated space—the measurement technique is explained below (Fig. 10.1). The other main alternative is to percutaneously fix pins to the different bones, let these protrude from the skin, attach markers for a video-based movement analysis system to the pins, and proceed to do the measurements while the initial local anesthetic is still effective. Most of the bones of the ankle and foot are easily accessible to both methods, although the talus, for example, offers some challenges. Specific anatomical and physiological considerations for using bone anchored markers are listed in the appendix. To summarize, the possible means of examination available to capture the true skeletal motion of the foot are: (1) RSA; (2) video-based analysis of kinematics of markers rigidly attached to bone; (3) video-based analysis of kinematics of rigid bodies (bones) captured from fluoroscopy, ultrasound, or other nonskin-penetrating methods and based on previous knowledge of geometry or other important properties of the involved bodies.

FIGURE 10.1 Foot with radio opaque markers [5].

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FIGURE 10.2 Radiostereometric analysis setup with calibration and experimental situations.

Alternative 3 was, already in the 1980s, considered attractive, and some of the studies described in this text were performed to improve the process of validating such methods. Note that techniques in Alternative 3 do not require invasive markers or bone pins.

10.3

Radiostereometric analysis

10.3.1 Technique RSA [6] traditionally means: (1) Inserting at least three markers rigidly fixed to each segment (bone) to be studied (Fig. 10.1). (2) Obtaining two simultaneous X-ray images in a calibrated volume (using markers in known positions in a calibration device) of the segments to be studied in a neutral position (reference examination) (Fig. 10.2). (3) Obtaining similar paired X-rays, also in a similarly calibrated space of the segments under study in any other situation that is of relevance (e.g., different foot position or later in time when studying changes after surgery). (4) Calculating the relative movements of the segments between examinations. Apart from the issue of radiation exposure, only point 1 differs from modern video-based kinematic analysis. RSA provides excellent accuracy of measurement, but there has historically been a lack of equipment for use in a dynamic setting. More recently, methods allowing dynamic analysis have been introduced [7]. Although RSA may mainly have become known as a method for determining implant stability in hip and knee arthroplasty, early uses of RSA included foot and ankle biomechanics such as study of fracture fixation stability in addition to kinematical studies described below [8]. The main body of RSA study into foot and ankle kinematics was performed at the Karolinska Institute, Stockholm, Sweden in the late 1980s. The main limitations were the lack of dynamic assessment as stated, plus the limited number of individuals that were studied. In cadaveric experiments, the input motion has usually been a controlled movement of one bone, with measured resulting movements of other bones as the outcome [9]. With in vivo studies, positions such as standing on a sloping platform or standing on a heel block have been referenced to neutral stance. The studies described below, unless otherwise stated, were performed on eight healthy volunteers and were approved by the local ethics committee. The joint kinematics of several individual foot and ankle joints were explored.

10.3.2 Talocrural joint The main controversy regarding the talocrural joint in the 1980s related to whether there was a single axis of motion (a hinge joint) as suggested by Inman [10] and others, or a more complex pattern such as that described by Hicks [4] and Barnett and Napier [11], indicating different joint axis orientations for plantarflexion and dorsiflexion. Over time, both hypotheses have gained support in different studies. The in vivo RSA studies of eight participants carried out at the Karolinska Institute by Lundberg et al. [12] indeed indicated that both patterns may be possible (Fig. 10.3). Some investigators had noted that the talocrural joint had a range of transverse plane rotation that was not negligible [13], in particular in light of the relatively modest range of plantarflexion/dorsiflexion used in walking, but the possibility of this influencing the perceived joint axis orientation during complex movements had not been discussed to any major degree.

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FIGURE 10.3 Individual axes for plantarflexion and dorsiflexion in eight examined participants.

(A)

(B)

Internal rotation External rotation Pronation Supination Plantarflexion Dorsiflexion

Internal rotation External rotation Pronation Supination Plantarflexion Dorsiflexion

FIGURE 10.4 Transverse plane projections for movement axes of the talocalcaneal (left) and talonavicular (right) joints, averaged over eight participants. Editor’s note: this reproduced figure includes the terms “pronation” and “supination” which in the context of the current chapter and book refer to “eversion” and “inversion.”

10.3.3 Relationship between the joints distal to the talus Traditionally the concept of a “subtalar” joint had at most represented the anterior and posterior talocalcaneal joints. The importance of the interaction between the talocalcaneal, talonavicular, and calcaneocuboidal joints was noted by researchers in the Netherlands as a “closed kinematic chain” [14]. The RSA studies at the Karolinska Institute were not designed to study this particular interaction, but would seem to confirm the importance of the talonavicular joint in the kinematics of the ankle and foot. The relative orientations of the axes of rotation of the talocalcaneal and talonavicular joints (the calcaneocuboidal joint was not studied at that time) were determined in the Karolinska Institute studies (Fig. 10.4). What may seem obvious to anyone looking at these bones and joints today (i.e., that the talocalcaneal joint will indeed act much as a hinge whereas the talonavicular joint is physiologically closer to a “ball and socket” joint) was not well understood at the time [15].

10.3.4 Transferral of rotation between the leg and the foot Levens et al. [1] and others had noticed the fact that in the stance phase of gait, the pelvis will rotate, which will force the femur to rotate, which in turn will force the tibia to rotate, while the foot will not rotate around a vertical axis (i.e., an axis normal to the transverse plane). The most well-known illustration of this may still be Inman’s oblique hinge [10]. According to Karolinska Institute studies, the rotation transferral mechanism between internal/external rotation of the leg and eversion/ inversion of the foot was much more obvious in the range from perceived neutral rotational position of the leg in relation to the foot into external leg rotation than in the range of internal leg rotation [16,17]. In fact, the kinematic pattern of the

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proximal joints is opposite to that of the distal joints of the arch (Fig. 10.5). This would be expected to be caused by the examined participants keeping the foot “as plantigrade as possible,” thus allowing the forefoot to counteract the movement induced from the hindfoot, which would otherwise have led them to standing on the outer edge of the foot. While this means that something that would be controlled in a cadaveric study in uncontrolled, it would certainly mimic “real life,” where the foot can be expected to be the site of such compensatory movements all the time.

10.3.5 Ankle mortise width This had been a long-term controversy, with some experts claiming that dorsiflexion of the talocrural joint would cause a separation of the tibia and the fibula due to the perceived larger width of the anterior part of the talar dome, while others, such as Inman [10], claimed that the shape of the talar dome would not be of importance. The contested importance mainly concerned the situation after rigid fixation of the mortise width after an ankle fracture. The Karolinska Institute studies indicated that there would indeed be a change in mortise width during a plantarflexion/dorsiflexion movement, but that the major part of this change would occur between a neutral position and plantarflexion, thus seemingly confirming the idea that this factor was unimportant in ankle fracture surgery [18] (Fig. 10.6).

10.4 Applications and significance of studies using intracortical pins for foot and ankle kinematics The intracortical pin method was seen as a means of avoiding the limitations of the RSA method to static or quasidynamic conditions. Reflective markers attached to metal pins firmly anchored in the relevant bone segments facilitated kinematic analysis during normal motion. Similar to the batch of RSA studies primarily underlying the above research, the series of studies using intracortical pins for the investigation of ankle and intrinsic foot kinematics revolved primarily around the Karolinska Institute in Sweden. Levens et al. [1] were the first known researchers to implement intracortical pins in human research in an investigation of lower limb kinematics. In more recent times, the first application of bone anchored pins with markers for motion analysis were investigations of the kinematics of horse motion looking primarily at skin displacement of the horse’s limbs during locomotion in 1987 and 1988 [19 21]. Human applications were initiated in the early 1990s in the PhD work by Irene McClay investigating tibial, femoral and patellar kinematics in participants with patellofemoral pain [22]. The first studies performed at Karolinska Institute were pioneering studies on knee kinematics [23,24], while another Swedish group also used the method in investigations of knee kinematics [25].

10.4.1 Skin movement artifact in foot and ankle kinematics A central possibility recognized in the intracortical kinematics research by the foot and ankle biomechanics community was the intuitive idea that skin marker accuracy could be validated during dynamic motion, as demonstrated in, at that time recent, equine limb and human knee research. An inherent problem with skin marker based kinematic analysis, and the intuitive basis for the development of invasive methods, is the error introduced when a skeletal segment is approximated by the skin covering it. An example of a pioneering study in this respect is Reinschmidt et al. [23], who Sup 15 (deg)

Tal–Tib Nav–Tal

10 Cal–Tal Cun–Nav 5 Met–Cun 0

Pro

5

20 Internal

10

0 Input

10 External

FIGURE 10.5 Resulting sagittal axis rotations from input internal/external leg rotation (average for eight participants). Tal-Tib 5 talus relative to tibia; Nav-Tal 5 navicular relative to talus; Cal-Tal 5 calcaneus relative to talus; Cun-Nav 5 medial cuneiform relative to the navicular; MetCun 5 first metatarsal relative to medial cuneiform.

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1.5

mm

Widening 1.0

0.5

0

– 0.5

– 1.0 Narrowing

degrees – 1.5 30 20 10 Plantarflexion

0

10 20 30 Dorsiflexion

FIGURE 10.6 Mortise width (measured as fibula medial/lateral translation relative to tibia) in relation to plantarflexion/dorsiflexion position (average for eight participants) [18].

analyzed differences in tibiocalcaneal rotations and translations from external marker based and intracortical pin based kinematics during running. The data indicated that the shape of curves presenting joint rotations were similar between the two methods but that there was a general overrepresentation of rotation magnitudes when using external markers. For example ankle eversion was measured as 16 versus 8.6 degrees respectively. It should be noted that these data were based on external markers attached to the running shoes rather than the skin directly. The potential and importance of such information was obvious, and a subsequent study of ankle joint kinematics utilized the same initiative by conducting motion analysis with simultaneous skin and intracortical marker sets for the calcaneus, talus, tibia, and fibula [26]. Westblad et al. calculated root mean square (RMS) differences between skin and intracortical pin marker tibiocalcaneal and talocalcaneal rotations and found no systematic over- or underrepresentation of skin marker data rotation range of motion in any plane. The RMS differences reported for tibiocalcaneal rotations were 2.5, 1.7, and 2.8 degrees for inversion/eversion, plantarflexion/dorsiflexion and abduction/adduction respectively and an RMS difference 2.1 degrees was shown in the determination of talocalcaneal inversion/eversion. Advances in the pin and marker setup constructions, in particular the introduction of smaller pins and marker arrays, permitted the insertion into considerably more bone segments and raised the question of validating the many advanced multisegment foot models that had been proposed in the literature. In conjunction with these developments, the optoelectronic systems were also improved and the resolution became sufficient to identify smaller markers in the calibrated volume. Nester et al. [27] defined leg, calcaneus, navicular/cuboid, medial forefoot, and lateral forefoot segments with a skin marker model in a validation study. They used both individual markers attached to the skin and several markers on rigid plates attached to the skin. The defined segments were compared to tibia, calcaneus, navicular, cuboid, and first metatarsal segments represented by intracortical pins. A necessary limitation with this more complete representation of the intrinsic foot segments was that the bone pin configuration did not permit simultaneous application of skin markers purely due to the skin surface area available (except for the tibia and calcaneus). Skin and intracortical pin trials were therefore conducted separately. Nester et al. again found no systematic pattern between the kinematics calculated from external or intracortical markers and concluded that it was unlikely that any one rigid body foot model defined by external markers would be preferable to another in terms of accurate representation of the foot and ankle bones [27]. A common criticism of the intracortical pin method is the doubt raised concerning how natural gait can be with multiple metal pins drilled into foot and ankle segments and that these sites are to some extent also anesthetized. This is a clearly justified concern for all these studies, maybe in particular in studies such as Nester et al. [27], where trials with and without intracortical pins needed to be compared. Several studies (e.g., Arndt et al. [28]) therefore analyzed for example ground reaction forces (GRF) during the stance phase of walking with and without intracortical pins and Maiwald et al. (Fig. 10.7) conducted a comprehensive analysis of 3D kinematics, GRF, and plantar pressure during walking and concluded that, despite some variation between participants, there were no intra-subject systematic effects of the intracortical pins on any of the variables analyzed [29].

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FIGURE 10.7 Location of the intracortical pins for a representative foot [29].

10.4.2 Ankle kinematics The foot and ankle RSA work by Lundberg et al. [11,14] and also similar research on the knee by van den Bogert et al. [24] raised the question of the position and orientation of the tibiocalcaneal and talocalcaneal axes during dynamic motion and in particular walking. Using the same dataset as in Westblad et al. [26], Arndt et al. [28] described the inclination and deviation of the instantaneous helical axes of rotation. Once again, the differences between participants were large in the helical axis orientation in both joints. As expected, Arndt et al. reported that while the majority of the overall ankle inversion/eversion occurred at the talocalcaneal joint (range of 4.6 8.3 degrees), the contribution of the tibiotalar joint cannot be ignored (range of 4.5 6.3 degrees) [28].

10.4.3 Foot and ankle basic research in walking and running kinematics Of paramount interest in the foot biomechanics research community was the accurate description of complex intrinsic foot kinematics to provide fundamental data in the understanding of gait mechanics, clinical treatment of foot disorders, podiatric interventions, and also footwear technology. This had previously not been possible, but it was facilitated by the successful innovation of the intracortical pin methodology. A requirement for seminal work with a basic research focus of purely descriptive nature was identified to provide accurate information for application in these research fields. Lundgren et al. described comparative motion of the fibula, tibia, talus, calcaneus, navicular, cuboid, medial cuneiform, first metatarsal and fifth metatarsal during walking both in terms of ROM and also as time series of motion (Fig. 10.8) [30]. An obvious extension addressing the same dearth in basic research data as seen in walking, was identified in running research. Tibiocalcaneal kinematics during running were investigated using film based motion analysis by Stacoff et al. [31]. They compared barefoot and shod running (including numerous different shoes and orthotic inserts). Only small and nonsystematic differences were seen between conditions (less than 2 degrees) especially when compared to the inter-subject differences reported (as large as 10 degrees). Only the most extreme conditions showed any consistent differences. It was concluded that tibiocalcaneal kinematics did not differ significantly between barefoot and shod running and that previous differences reported in the literature may have been due to the use of skin or shoe-mounted markers. The more comprehensive data provided by the advanced methodology with a greater number of inserted pins combined with the increased precision and ease of analysis given by the optoelectric motion capture system described above, also permitted the establishment of a unique dataset describing slow running [32]. Only slow running could be performed due to safety reasons concerning the behavior of the pins during more dynamic activity and also simply because the laboratory runway (9.5 m) was not sufficiently long to allow a steady state faster running before needing to decelerate. Running data could subsequently be compared to walking data and fundamental differences described. A comparison of ranges of motion occurring about specific joints is provided (Fig. 10.9A). The ROM was greater during walking than running at each of the joints analyzed (Fig. 10.9B), indicating that the intrinsic muscular function in the foot provides a stiffer system during running than walking. An attempt to define functional units in the human foot and provide a meaningful basis for subdivisions of the entire foot during gait analysis was presented by Wolf et al. [33]. Angle-angle diagrams were used to identify groups of segments moving as functional units. Two functional units were identified: the medial ray from the navicular to the first metatarsal and the combined navicular and cuboid, which is important information for justifying simplified foot models.

10.4.4 Applied studies of orthoses and shoe conditions Despite the plethora of anatomical-biomechanical scientific questions that could be explored with intracortical foot and ankle kinematics, there is also an endless spectrum of applied clinical and performance applications that could be investigated. These are obviously limited by the invasive nature of the methodology, which restricts the number of participants, limits the use of shoes or other devices that may interfere with the pins and at least at this stage has precluded patients with relevant injuries or diseases from participating. These issues are important and need to be taken into

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FIGURE 10.8 Computed tomography (CT) scan demonstrating the intracortical pins for a representative foot. (A) Anterior view. (B) Posteriorlateral view [30].

account when designing studies and also by the relevant ethical committees who need to determine the integrity and safety of the participants in these studies in relation to the risk/benefit weighting in terms of scientific knowledge gains. Some studies of importance have, however, been attempted. For example Stacoff et al. [34] and Liu et al. [35] addressed the controversial issue of the efficacy of plantar orthoses in adjusting intrinsic foot skeletal alignment. By facilitating a direct measurement of the influence of orthoses on intersegmental alignment and rotations the first data could be presented describing for example effects upon the subtalar joint. In the study by Stacoff et al. [34] it was concluded that a medially placed foot orthosis did not modify tibiocalcaneal kinematics during the stance phase of running. Liu et al. [35] found that tibiotalar and talocalcaneal kinematic responses to “anti-pronation” (i.e., anti-eversion) orthoses contradicted the existing paradigms in that the changes did not only occur at the talocalcaneal joint. This supported the data from Arndt et al. [32] that eversion and inversion of the ankle joint are not restricted to the talocalcaneal joint. Stacoff et al. [36] investigated the effects of shoe sole construction on tibiocalcaneal kinematics during the stance phase of running and found no clear relationship, indicating that the kinematics may be individually unique and difficult to systematically modify with shoe modifications. They also reported a clear overrepresentation of rotations when measured with skin and shoe markers as opposed to intracortical pin markers. Similarly, Arndt et al. [37] found no direct transferral of decreased bending stiffness of a shoe sole on the intrinsic foot kinematics during walking.

10.5

Limitations and future directions

Intracortical pins may affect walking, but Maiwald et al. [29] analyzed the effects on gait occurring due to inserted pins and found that local anesthesia and the presence of bone pins still allow a valid gait pattern to be analyzed. Modern dynamic X-ray imaging techniques have also indicated that intracortical pin data may lead to apparent collisions between the bones, which may be due to some oscillation of the protruding pins to which the marker arrays are attached.

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(A)

Mean (SD)

Mean (SD)

Mean (SD)

Mean (SD)

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7.2 (2.4)

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5.3 (2.0)

13.3 (1.4)

frontal

11.3 (3.9)

8.8 (4.4)

10.4 (6.3)

5.4 (1.0)

10.4 (3.7)

8.1 (2.0)

8.9 (4.3)

6.2 (4.2)

6.1 (1.1)

9.8 (2.1)

Mean (SD)

Mean (SD)

Mean (SD)

transversal

(B)

Mean (SD) sagittal

7.8 (1.2)

5.5 (2.0)

7.1 (4.0)

4.9 (2.5)

11.4 (1.6)

frontal

6.3 (1.3)

5.5 (2.0)

8.1 (2.6)

5.3 (1.9)

5.1 (0.6)

transversal

6.9 (3.3)

5.6 (1.8)

4.1 (1.1)

4.3 (1.4)

9.6 (2.4)

FIGURE 10.9 (A) Ranges of motion in selected joints during walking. (B) Ranges of motion in the same joints during slow running. Courtesy of Peter Wolf.

Despite the important expansion of the knowledge base concerning foot and ankle biomechanics that has been achieved by the era of intracortical pin studies, the limitations and possible problems are evident. As with any invasive technique there are risks of infection, damage to adjacent anatomical structures, allergic reactions to antibiotics, and pain. These have been successfully avoided or minimized in the Karolinska Institute studies and the positive experience gained has also facilitated easier ethical approval. However, methods for accurately determining foot and ankle kinematics noninvasively would be much preferable if they were available. Recent advances in high speed double X-ray video techniques have provided promising results (e.g., Welte et al. [38]). These techniques use pattern recognition of the skeletal segments based on computer tomography images, which eliminates the intermediate step of registering bones to markers, and markers to motion because the bone motion is acquired as part of the process. Although promising there are still some problem areas with such techniques such as a very small target volume for data recording, extensive processing time, and radiation exposure. The largely unfunded nature of the research conducted at Karolinska Institute, based to a large extent upon the enthusiastic interest of the researchers involved and the establishment of a unique and successful method, does limit the amount of data that can be processed. The extensive amount of data could be used for a considerable amount of new scientific questions but due to limitations in time available for this project, many still remain unanswered. The high international interest in collaborating is an indication of many biomechanists having such questions and wanting to initiate combined research projects, a sign of the possibilities of international inquisitive science with innovative and excited researchers.

Appendix:

Insertion of markers in bones of the foot and ankle

Inserting X-ray markers: The main concern relating to markers for RSA is that there should be at least three per bone as widely distributed as possible. In particular, the second requirement may be difficult to meet in some of the smaller bones of the foot. The

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analysis software typically provides quality measurements to help decide whether the distribution is good enough for high quality kinematic calculations. Markers are normally spherical, 0.5 1.0 mm in diameter and inserted through a 1.0 1.5 mm needle with a device for pressing the marker firmly into the bone. The needle is drilled into the bone with reciprocating motion (not rotated) which obviates an outer tube for tissue protection. Concerns for individual bones: 1. Tibia. Easy to get a good marker distribution. Obviously, the neurovascular bundle behind the medial malleolus should be avoided. 2. Fibula: Easy to reach but important to get a maximum anterior/posterior and medial/lateral distance between markers. 3. Talus: The talus presents several challenges. One is that the talar neck (where the majority of markers normally are place) can be rather narrow. Another is to get one or two markers in from behind without interfering with the medial neurovascular bundle or the sural nerve. 4. Calcaneus: This is the easiest bone of the foot/ankle complex to put markers in. Not much cortical bone and numerous safe approaches from medial, lateral and posterior directions, with only the sural nerve to be avoided. 5. Navicular: Can be reached from the medial side and/or from above. The challenge is to avoid markers being placed more or less on a horizontal line from medial to lateral, attention has to be paid to getting markers near both the dorsal and the plantar aspect of the bone. 6. Cuboid, cuneiforms, metatarsals, and phalanges: Easy to reach. The dimension of the needle carrying the marker may be too large for toes; also, this will be an area where a satisfactory marker distribution may be difficult to obtain. Basic method for inserting pins: The pins used are in most cases pins normally used for external fixation of fractures. These are sufficiently stiff to prevent unnoticed tissue impingement, and they fulfill all criteria for use as temporary implants. The diameter is usually between 1.8 and 3.5 mm, depending on the dimensions of the target bone. The insertion site is ideally chosen as the point where relative movement between bone and overlying skin is minimal, but the risk of interference between the pin in question and other pins or anatomical features (such as the other foot or leg in walking or running studies) also has to be considered. Concerns for inserting pins in individual bones: 1. Tibia: The tibia is easy to reach, and the soft tissue coverage is relatively thin. There is no reason to go behind the medial malleolus, the metaphyseal area between the malleolus and the anterior tibial tendon readily presents itself as long as you keep the pin close to the sagittal plane. However, the tibialis anterior tendon needs to be avoided. The cutaneous branches of the peroneal nerve should not be a problem as long as you use a drill sleeve. Going in from a medial position should be avoided as it can lead to interference with the other leg/foot in gait; the pin should be oriented as much in the sagittal plane (i.e., from anterior to posterior) as possible. 2. Fibula: Easy to reach. The pin should enter from anterior/superior (and lateral) downward to avoid interference with the talar pin. 3. Talus: As for X-ray markers, the talus presents several challenges. One, again, is that the talar neck can be rather narrow, the main other concern is soft tissue interaction. The safe zone lies between the tibialis anterior tendon and the deltoid ligament. 4. Calcaneus: This, as with X-ray markers, is the easiest bone of the foot/ankle complex to insert bone pins into. It is easy to reach from the lateral or posterior direction or from the medial side if interaction with the opposite limb is not an issue. 5. Navicular: This is where fluoroscopy guidance is of vital importance. The shape of the navicular means that a bone pin needs to go in from above and rather straight down to avoid the risk of interference with the joints proximal and distal to the navicular. 6. Medial cuneiform, cuboid, first and fifth metatarsals, first proximal phalanx: These bones are easy to reach, and there is minimal risk of soft tissue snagging. The risk of interference with the other foot or pins in adjoining bones should be the only concerns. 7. Other cuneiforms, metatarsals, and phalanges: The main concerns are extensor tendons and the risk of interference with pins in adjoining bones. In the lesser phalanges, it would also be difficult to insert a pin stiff enough to carry a marker array without creating an increased risk for either movement artifact or fracture.

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References [1] [2] [3] [4] [5] [6] [7] [8] [9] [10] [11] [12] [13] [14] [15] [16] [17] [18] [19] [20] [21] [22] [23] [24] [25] [26] [27] [28] [29] [30] [31] [32] [33] [34]

Levens A, Inman V, Blosser J. Transverse rotation of the segments of the lower extremity in locomotion. J Bone Jt Surg 1948;30-A:859 72. Close J, Inman V, Poor P, Todd F. The function of the subtalar joint. Clin Orthop 1967;50:159 79. Inman V. The influence of the foot-ankle complex on the proximal skeletal structures. Artif limbs 1969;13:59 65. Hicks J. The mechanics of the foot. I. The joints. J Anat 1953;87(4):345 57. Lundberg A, Goldie I, Kalin Bo, Selvik G. Kinematics of the ankle/foot complex: plantarflexion and dorsiflexion. Foot Ank Int 1989;9 (4):194 200. Selvik G. Roentgen stereophotogrammetry. A method for the study of the kinematics of the skeletal system. Acta Orthop Scand 1989;60(Suppl. 232):1 51. Gustafson J, Elias J, Fitzgerald G, Tashman S, Debski R, Farrokhi S. Combining advanced computational and imaging techniques as a quantitative tool to estimate patellofemoral joint stress during downhill gait: a feasibility study. Gait Posture 2020;84:31 7. Ka¨rrholm J. Roentgen stereophotogrammetry. Review of orthopedic applications. Acta Orthop Scand 1989;60(4):491 503. van Langelaan E. A kinematical analysis of the tarsal joints. Acta Orthop Scand 1983;54(Suppl 204):1 269. Inman V. The joints of the ankle. Baltimore: Williams & Wilkins; 1976. p. 1976. Barnett C, Napier J. The axis of rotation at the ankle joint in man. Its influence upon the form of the talus and the mobility of the fibula. J Anat (Lond) 1952;86:1 9. Lundberg A, Svensson OK, Ne´meth G, Selvik G. The axis of rotation of the ankle joint. J Bone Jt Surg 1989;71-B(1):94 9. Rasmussen O, Tovborg-Jensen I. Mobility of the ankle joint. Acta Orthop Scand 1982;53:155 60. Huson A. Perspectives in human-joint kinematics. In: Huiskes R, Van Campen D, De Wijn, editors. Biomechanics: principles and applications. The Hague: Martinus Nijhoff Publishers; 1982. p. 31 45. Lundberg A. Kinematics of the ankle and foot. In vivo roentgen stereophotogrammetry. Acta Orthop Scand 1989;233(Suppl. 1):1 24. Lundberg A, Svensson OK, Bylund C, Goldie I, Selvik G. Kinematics of the ankle/foot complex part 2: pronation and supination. Foot Ankle 1989;9(5):248 53. Lundberg A, Svensson OK, Bylund C, Selvik G. Kinematics of the ankle/foot complex part 3: influence of leg rotation. Foot Ankle 1989;9 (6):304 9. Svensson OK, Lundberg A, Walheirn G, Selvik G. In vivo fibular motions during various movements of the ankle. Clin Biomech 1989;4 (3):155 60. van den Bogert AJ, van Weeren PR, Schamhardt HC. Correction for skin displacement errors in movement analysis of the horse. J Biomech 1990;23:97 101. van Weeren PR, van den Bogert AJ, Barneveld A. A quantitative analysis of skin displacement in the trotting horse. Equine Vet J 1990;(Suppl. 9):101 9. van Weeren PR, van den Bogert AJ, Barneveld A. Quantification of skin displacement in the proximal parts of the limbs of the walking horse. Equine Vet J 1990;(Suppl. 9):110 18. McClay I, Cavanagh P, Kalenak A, Sommer H. 3-dimensional kinematics of the patellofemoral joint during running. Proc Am Soc Biomech 1991;15:146 7. Reinschmidt C, van den Bogert A, Murphy N, Lundberg A, Nigg B. Tibiocalcaneal motion during running, measured with external and bone markers. Clin Biomech 1997;12(1):8 16. van den Bogert AJ, Reinschmidt C, Lundberg A. Helical axes of skeletal knee joint motion during running. J Biomech 2008;41(8):1632 8. Ramsey D, Wretenberg P, Benoit D, Lamontagne M, Ne´meth G. Methodological concerns using intra-cortical pins to measure tibiofemoral kinematics. Knee Surg Sports Traumatol Arthrosc 2003;11:344 9. Westblad P, Hashimoto T, Winson I, Lundberg A, Arndt A. Differences in ankle-joint complex motion during the stance phase of walking as measured by superficial or bone anchored markers. Foot Ankle Int 2002;23(9):856 63. Nester C, Jones R, Liu A, Howard D, Lundberg A, Arndt A, et al. Foot kinematics during walking measured using bone and surface mounted markers. J Biomech 2007;40(15):3412 23. Arndt A, Westblad P, Winson I, Hashimoto T, Lundberg A. Ankle and subtalar kinematics measured with intracortical pins during the stance phase of walking. Foot Ankle Int 2004;25(5):357 64. Maiwald C, Arndt A, Nester C, Jones R, Lundberg, Wolf P. The effect of intracortical bone pin application on kinetics and tibiocalcaneal kinematics of walking gait. Gait Posture 2017;52:129 34. Lundgren P, Nester C, Liu A, Arndt A, Jones R, Stacoff A, et al. Invasive, in vivo measurement of rear, mid and forefoot motion during walking. Gait Posture 2008;28(1):93 100. Stacoff A, Nigg B, Reinschmidt C, van den Bogert A, Lundberg A. Tibiocalcaneal kinematics of barefoot vs shod running. J Biomech 2000 (a);33:1387 95. Arndt A, Wolf P, Liu A, Nester C, Stacoff A, Jones R, et al. Intrinsic foot kinematics measured in vivo during the stance phase of slow running. J Biomech 2007;40(12):2672 8. Wolf P, Stacoff A, Liu A, Nester C, Arndt A, Lundberg A, et al. Functional units of the human foot. Gait Posture 2008;28(3):434 41. Stacoff A, Reinschmidt C, Nigg B, van den Bogert A, Lundberg A, Denoth J, et al. Effects of foot orthoses on skeletal motion during running. Clin Biomech 2000;15:54 64.

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[35] Liu A, Nester C, Jones R, Lundgren P, Lundberg A, Arndt A, et al. The effect of an anti pronation foot orthosis on ankle and subtalar kinematics. Med Sci Sports Exerc 2012;44(12):2384 91. [36] Stacoff A, Reinschmidt C, Nigg B, van den Bogert AJ, Lundberg A, Denoth J, et al. Effects of shoe sole construction on skeletal motion during running. Med Sci Sports Exerc 2001;33(2):311 19. [37] Arndt A, Lundgren P, Liu A, Nester C, Maiwald C, Jones R, et al. The effect of a midfoot cut in the outer sole of a shoe on intrinsic foot kinematics during walking. Footwear Sci 2013;5(1):63 9. [38] Welte L, Kelly L, Kessler S, Lieberman D, D’Andrea S, Lichtwark G, et al. The extensibility of the plantar fascia influences the windlass mechanism during human running. Proc R Soc B 2021;288:20202095. Available from: https://doi.org/10.1098/rspb.2020.2095.

Chapter 11

Biplane Fluoroscopy Eric Thorhauer1,2 and William R. Ledoux1,2,3 1

RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 2Department of

Mechanical Engineering, University of Washington, Seattle, WA, United States, 3Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Abstract In the time since Wilhelm Roentgen first accidently made X-ray images of his wife Anna’s hand, radiographic technologies have gradually evolved to allow for dynamic imaging of the foot and ankle. In this chapter, we discuss how biplane fluoroscopy systems have been developed and modified specifically for overcoming the obstacles of quantifying full dynamic foot and ankle biomechanics that are associated with motion capture or computed tomography technologies. Biplane systems can precisely resolve the motions of small foot bones, and they are commonly synchronized with force plates, electromyography systems, and motion capture as part of a multifaceted biomechanical assessment strategy. Biplane systems have improved from pairs of simple stock intact C-arm fluoroscopy units, to the modification of these stock systems, to fully custom fluoroscopy suites. The initial obstacles of developing and validating hardware and software as well as those involved in administering in vivo foot and ankle studies present exciting challenges and avenues of research that draw on contributors from a variety of technical backgrounds. This chapter provides a general overview of both the biplane hardware and software, and their application to a variety of clinical issues related to ambulation and revealing the intricate functions of the foot and ankle. The future uses of biplane systems include assessments of joint arthrokinematics, ligament deformations, and soft tissue strains to address both fundamental and complex questions related to foot function, which will serve to elucidate biomechanics and undoubtedly generate valuable insights for clinicians and researchers alike for decades to come.

11.1

Introduction

Biplane systems can be used to see inside the foot, revealing the motions of all the bones instead of grouping them into kinematic segments like traditional retroreflective motion capture requires. Furthermore, biplane fluoroscopy derives bone motion directly from the imaging instead of inferring it from skin-mounted markers that potentially introduce errors [1,2]. Compared to other medical imaging modalities like computed tomography (CT), magnetic resonance imaging (MRI) or ultrasound, biplane fluoroscopy has an adequate combination of capture volume size, temporal resolution, and spatial resolution for quantifying subtle differences in three-dimensional (3D) bone motion with minimal aliasing. Flexibility in the geometry of biplane fluoroscopy capture volumes permits testing of dynamic, functional tasks by removing obstacles to test subjects like gantry beds and scanning bores. The full power of biplane systems is leveraged when synchronized with other acquisition systems (electromyography, force plates, motion capture, etc.) to round out a more complete biomechanical analysis that can frame the biplane-derived data in the context of the subject’s wholebody kinematics and kinetics. Together, these tools paint a more complete picture of the biomechanical roles of the foot in our daily activities. However, three main obstacles need to be addressed to facilitate the widespread adoption of this technology. First, executing biplane fluoroscopic analyses requires a substantial infrastructure investment, including designing a specialized laboratory space, and the integration of complex, expensive X-ray generation and detection hardware. There are many infrastructure considerations related to the specific research question under investigation (e.g., a laboratory emphasizing full foot kinematics may require a configuration that allows for oblique imaging with a large image detector). Second, the bone-tracking process requires specialized software that generates voluminous amounts of data that Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00044-5 © 2023 Elsevier Inc. All rights reserved.

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must be effectively managed for studies with large cohorts or longitudinal aspects. The software solutions are often custom developed and require careful validation to ensure bias is minimized and precision maximized. Third, the ionizing radiation that is received from two fluoroscopy sources and a CT scan during biplane testing present risk to both subjects and lab personnel. However, by adhering to ALARA (as low as reasonably achievable) principles [3], biplane systems can ethically and effectively be used for longitudinal in vivo human biomechanical research studies. With proper planning, it is possible to minimize exposure to subjects and staff while simultaneously fulfilling the research needs of producing usable image data that balance and respect the risks involved in radiography. These risks are best addressed in the study design and protocols, but also in the design of the laboratory layout itself including but not limited to: X-ray shielding and aiming strategies, pre-planning system geometry in virtual biplane software, attempts at dose estimation for subjects, the deliberate and careful selection of X-ray imaging parameters, and the development of software solutions that move toward a reduction in total dose. Two hardware technologies that drastically reduce total dose are pulsed X-ray generators instead of continuous biplane sources, and the newest high-efficiency detector CT systems. A strategy of total dose reduction makes the best use of an IRB-approved dose allowance, and any reduction from that limit facilitates the collection of more trials of the subjects’ motion. More trials can better quantify the variation in motion for a given joint or body or assess a different task or a variable of interest (e.g., orthotic elevation or brace stiffness). While the doses subjects receive from biplane imaging equipment are well below the currently recognized thresholds for developing cancers based on conservative risk estimates [4,5], dose reduction is still motivated from an ethical standpoint as these exposures are for the purposes of research and not medical diagnostics. Continued advancements in hardware and software will jointly reduce long-term risks to subjects while addressing the challenges of tracking bones in noisy fluoroscopy images and segmenting bones from low-dose CT scans that are inherently noisier than traditional clinical scans. While the design, execution, and processing of biplane fluoroscopy foot and ankle studies present many technical challenges, the benefits of these efforts are deeply rewarding. Acting as a bridge between the research and clinical realms, biplane laboratories draw on the skills of engineers from a variety of backgrounds to address these obstacles. To effectively answer clinical research questions, the engineers need to foster cooperation and mutual respect with multiple types of clinicians that can aid study design focused around their subjects and ALARA principles. These technologies will continue to evolve, so this chapter cannot possibly cover all the intricacies of biplanar fluoroscopy. However, we hope it serves as a starting point to motivate collaboration between engineers and clinicians alike to broaden and deepen the impact of this technology on the field of foot and ankle biomechanics research.

11.2

Background and history of biplane fluoroscopy

11.2.1 Overview of how X-ray imaging works Exploitation of the attenuating properties of X-rays for medical imaging was pioneered shortly after Wilhelm Roentgen’s initial discovery in 1895. Depending on the composition of the tissue and the energetics of the incident beam, X-rays leaving the source undergo a variety of transport and interaction phenomena with the subject [6]. X-rays may be completely absorbed, attenuated, and transmitted on to the detector, or blocked or scattered into a new trajectory. Absorbed rays contribute to subject radiation dose, and subsequently risks associated with ionization radiation. Scattered rays can contribute to image noise. The degree of attenuation for X-rays passing through the subject onto the detector is influenced by the mass density properties of the tissues and the incident beam energetics. The increased attenuation through bony tissues relative to the surrounding soft tissues is what is exploited in radiography to form the familiar X-ray projections. Similar principles are at work in CT imaging, which uses many X-ray back projections from varying angles to reconstruct a full 3D volume of voxels (i.e., the volumetric equivalent to a pixel) each assigned a Hounsfield Unit related to the tissue’s X-ray attenuation properties relative to water. X-ray source parameters like tube voltage and current, and firing rate are adapted to the study of interest with the concurrent goals of maximizing the usable information in the biplane images while adhering to ALARA principles. These settings may be carefully selected prior to living subjects imaging via experimental testing on cadaveric surrogates simulating the activity of interest and adapted on a per subject basis.

11.2.2 Overview of biplane system history and evolution Biplane X-ray systems were developed in the 1930s [7], driven by the needs of surgeons performing advanced procedures, who greatly benefitted from the real-time imaging afforded by fluoroscopy and the extra spatial information

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obtained from a second orthogonal X-ray view. In the 1970s, Selvik and colleagues pioneered RSA or “Roentgen stereophotogrammetric analysis” for quantifying subtle migrations of implants in orthopedic patients [8 10]. The technology was next adapted to address orthopedic questions related to skeletal motion [11 13]. While the methods and techniques associated with biplane X-ray systems for surgical use are a great starting point for developing a powerful tool for biomechanics research, there are limitations and considerations that require extensive modification of original equipment manufacturer (OEM), medical-purpose biplanar and single-plane imaging systems, namely their configurability into arbitrary capture volumes and their sampling rate. As researchers pursue questions related to the dynamic behavior of the foot during tasks like gait or stair navigation, these biplanar imaging systems will continue gradual evolutions in both hardware and software to address early limitations in bone tracking. Although not directly applicable to the foot and ankle, it should be noted that single and biplane fluoroscopy systems have demonstrated sub-millimeter and sub-degree biases and precisions during bone-based tracking [12,14 16].

11.2.2.1 Intact C-arm systems for foot bone tracking Early efforts to implement fluoroscopy systems for quantifying foot motion utilized intact C-arm fluoroscopy units that are commonly used intraoperatively and in clinical settings [17,18]. Manufactured by long-established medical imaging technology companies, these systems mount the X-ray source and image detector on a fixed C-shaped gantry, and include accompanying software and hardware for acquiring and viewing the images. Their fixed source-to-detector distance and the presence of the C-shaped gantry structure impose physical constraints on the types of subjects, anatomy, and functional activities that may be assessed in such a system.

11.2.2.2 Disarticulated C-arm systems for foot bone tracking To overcome these constraints, researchers disarticulated the X-ray sources and detectors, mounting them on custom gantries or supports, opening new opportunities for the kinds of studies possible [19 21]. For example, prior to disarticulation of biplane systems, only quasi-static approximations of gait were possible. Subjects can now walk unimpeded through disarticulated systems.

11.2.2.3 Custom dedicated biplane hardware for foot bone tracking The next evolution of biplane systems was to implement the X-ray sources and detectors of catheterization and angiography fluoroscopy suites; these devices were often developed to study other parts of the body and subsequently adapted to quantify foot motion [22 24]. By eliminating the physical C-arm, and increasing the source-to-detector distance, more powerful sources are required to produce images with usable quality. However, the flexibility of these designs allowed for improvements in the hardware, including: larger detectors with sufficiently fast response times, more configurable X-ray sources and system geometry, pulsed fluoroscopy and varying sampling rates, and the use of high-speed cameras at increasing resolutions.

11.3

Other techniques for tracking foot bone kinematics

Although it is growing in popularity, biplane fluoroscopy is still somewhat of a niche technology. As of the latest data available (mid-2020), there were less than 50 biplane laboratories worldwide [25], and only about a dozen groups that have studied the foot and ankle (see below). Prior to the implementation of biplane fluoroscopy, foot bone motion has been tracked using near infra-red cameras and retro-reflective markers [26,27], but these techniques require grouping multiple bones together into simpler rigid bodies. Moreover, skin motion artifact is a major concern at the foot and ankle. Finally, attaching markers to skin requires that subjects walk either barefoot or in severely compromised shoes, which precludes the study of the effects of footwear or orthoses. Despite the growing number of studies using retroreflective foot models, this technique is ill-suited to study specific foot joints. Other technologies have been employed to quantify foot bone motion. Marker-based X-ray stereophotogrammetry has been used to evaluate the orientations of the bones of the foot in multiple static positions via implanted tantalum balls and a specialized apparatus [28 31]. Time-sequence MRI has also been used to explore foot bone orientation [32 35]. Both of these techniques are limited by the static nature of the data collected. Retro-reflective markers mounted on bone pins have also been used to quantify foot bone motion during walking and slow running [36 39], however, the invasiveness of this methodology limits its utility to the research lab and precludes clinical studies. Moreover, none of these techniques work with footwear. In summary, current technologies to track foot bone motion,

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including: surface markers, marker-based X-ray stereophotogrammetry, time-sequence MRI, or bone pins, are ill-suited to measure barefoot and shod foot bone kinematics. Biplane fluoroscopy can track 3D joint positions accurately and avoid the inherent errors of marker placement and skin motion artifact present in retro-reflective motion analysis. Single plane (i.e., two-dimensional (2D)) fluoroscopy has been used to study heel pad deformation [40], hindfoot kinematics [41], and the effects of various immobilization techniques on ankle motion [42]. 2D fluoroscopy techniques work well in-plane, but these are not as accurate out-ofplane. And while the simplicity, flexibility, and utility of using a single fluoroscope are obvious and attractive, the complex anatomy of the foot—with many small, overlapping bones—and the importance of 3D motion, requires a biplane system. We mainly review the biplane systems that have been employed to quantify foot bone motion; systems used to study other parts of the body will not be discussed.

11.4

Challenges specific to foot and ankle tracking with biplane fluoroscopy

To quantify foot bone motion in an assortment of tasks, a biplane system must be configurable into capture volumes that do not impede or alter natural subject motion and, preferably, avoid or minimalize the beams directly hitting the more radiation sensitive anatomical regions. Allowing for six degree-of-freedom positioning of both the X-ray sources and detectors, however, requires a larger testing space for the equipment and protocols for enforcing proper beam-detector alignment. Subject-specific planning in aiming the biplane system can improve the usability of the fluoroscopy images and reduce subject radiation dose. Characterizing the motion of each bone in the foot during a full gait cycle in a single trial is only possible when the X-ray image detector is large enough to capture human adult anatomy while also allowing for the small deviations in foot placement from the center of the biplane system capture volume, and fast enough to image moving subjects without blur. Early C-arm based fluoroscopy systems have maximum detector sizes (31 cm) that necessitate the capture of multiple trials specifically targeting the forefoot or hindfoot in isolation, and maximum frame rates (B15 30 Hz) that limit assessments to slow or quasi-static activities. Additionally, the X-ray sources must be precisely and consistently controlled to start and stop firing at the appropriate instants to minimize subject dose while still ensuring the motion of interest was captured. This triggering can be initiated by force plate-based gait events (e.g., heel strike or foot placement), optoelectrical switches aimed across the subjects’ path of progression, or manually by the system operator. In many studies of the ankle and hindfoot, oblique sagittal views of the foot are adequate. This greatly simplifies the laboratory setup; all the components can be placed in the same plane as the floor. If the study design warrants imaging the forefoot or the foot in its entirety, the necessary hardware is more involved. A custom walkway traversing a biplane pit (Fig. 11.1) has been designed at the Center for Limb Loss and MoBility (CLiMB) at the Veterans Affairs Puget Sound Health Care System that is configurable into multiple walkway setups for assessing linear gait, turning tasks, stair negotiation, and more. A panel in the target volume portion of the walkway is made from radiotranslucent carbon fiber that does not appear on the fluoroscopic images. The corners of the panels are mounted with load cells to capture the ground reaction forces. Custom supports hold the X-ray detectors under the walkway as close to the foot as possible and an overhead structure supports the sources without impeding subject motion.

11.5

Overview of biplane hardware

The imaging chain of a dedicated custom biplane system consists of an X-ray source and an image intensifier paired with a high-speed digital camera (Fig. 11.2). X-ray image intensifier technology was introduced at the end of the 1940s, and despite some technical considerations overviewed below, works adequately for this application. Modern flat-panel detectors (FPD) have a high quantum detector efficiency that reduce the dose required to form a useable image without spatial distortions [43 45]. While this technology will continue to evolve, presently, FPDs that are large enough area to image the entire foot are only offered with sampling rates that are inadequate for capturing gait. Consequentially, this section will only detail systems based on the proven coupling of X-ray image intensifiers and high-speed video cameras that permit blur-free imaging of dynamic whole foot function. X-ray sources on custom biplane systems can be fully independently controlled in terms of aiming and beam parameters like current (mA), tube potential (kV), continuous versus pulsed beam firing, pulse width/exposure time, collimation, and internal filtering. The latest X-ray generators create precisely synchronized pulses from both biplane sources while the high-speed cameras are triggered at the optimal interval to capture the fluoroscopy images leaving the image intensifiers without appreciable motion blur or detector lag. Collimators allow for shaping of the primary beams to

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FIGURE 11.1 An overview of the biplane process from study inception and design during the planning phase, through the acquisition phase utilizing the biplanar fluoroscopy and computed tomography imaging systems, and finally the processing phase of generating useful biomechanical data from the image data. Careful decision-making during study planning and quality assurance measures during the acquisition phase save considerable time and effort at the complex stages of data processing.

reduce scatter noise incident on the other detector and avoid exposure to anatomical regions of subjects that are not under investigation. Internal filters impede low-energy portions of the polychromatic beam spectrum not strong enough to penetrate the subject and reach the detectors to form an image. This filtering of the beam, also known as “hardening,” also increases contrast between tissues by effectively increasing the mean beam potential by omission of these weaker components. Exiting the sources and passing the filtration and collimators, X-rays proceed toward the image intensifier assembly losing intensity following the inverse square law. At the image intensifier face they are filtered again by anti-scatter grids that act to block incoming scattered rays from oblique angles that contribute to image noise. These grids are optimized for fixed source-to-detector distances. The passing portions of the signal are next absorbed by the image intensifier input screen and converted first into light photons then electrons following a Poisson noise process. In the evacuated tube of the image intensifier body, internal magnetic fields guide the signal of electrons toward the anode of the output phosphor screen where they are converted into visible green light emitted out of the back of the image intensifier. By condensing the input signal of the image intensifier face (e.g., 30 40 cm diameter) into a smaller output signal (e.g., 7.5 cm diameter), a minification gain is achieved that amplifies the overall image brightness by several orders of magnitude. This output signal is redirected via a 90-degree lens (to reduce the physical depth requirements of the biplane pit beneath the walkway) with a high-speed cine grayscale video camera with a telephoto lens ( . 2 megapixels, .100 Hz, exposure time ,1.5 milliseconds). The electron trajectory within the evacuated image intensifier body is highly susceptible to interference from the presence of external electromagnetic fields (i.e., Earth’s field and lab equipment) and metal in structures supporting the biplane system and walkway. Additionally, the cine camera chain also introduces optical distortions as a function of the lens geometry, selected aperture, and focusing. This corruption of the fluoroscope images, manifest as spatial “S-curve” and pincushion distortion patterns, must be rectified in software to accurately reconstruct 3D position in a process known as “distortion correction” [46 48]. A distortion correction map is generated by imaging a known grid

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FIGURE 11.2 An overview of the Center for Limb Loss and MoBility (CLiMB) biplane system hardware, namely: (A) gantry-mounted pulsed X-ray sources, (B) walkway with an X-ray-transparent force plate, and (C) image intensifier and high-speed camera chain. Collimators and anti-scatter grids are utilized to reduce the exposed subject area and cross-scatter, and to improve image quality. The X-ray image intensifier converts the X-ray signal into a visible light signal that is directed to the sensor of a high-speed camera with an array of lenses.

object placed directly against the image intensifier input screen. Mathematical models are used to warp and interpolate the distorted raw biplane fluoroscope signals into usable spatially rectified images [49 56]. Additionally, the intensity of the fluoroscope image is not uniform in illumination due to the roll-off characteristics and fewer incident X-rays at the image intensifier periphery. To correct the pixel intensities, which are especially important to model-based bone tracking algorithms, a flatfield correction is required in software to compensate for the dimming image edges [57]. To reconstruct the geometry of the biplane system in software, a camera calibration process is required each time the system configuration is altered or re-aimed [58]. To perform the model-based tracking software procedure (outlined in the following section), volumetric subjectspecific bone models are needed from CT scans. Currently, highly efficient low-dose weightbearing CT scanners are available for laboratory and clinical application to the foot and ankle [59,60]. These bilateral foot scans are acquired in under two minutes and at a fraction of the radiation dose (typically , 4 mrem) of clinical scanners (typically 20 mrem). The dose received from clinical CT scanning typically represents the largest single contribution to the total received dose of biplane test subjects. By reducing the CT dose, more of an IRB-approved per-subject radiation dose budget may be used in biplane trials characterizing motion variation, other tasks, or a longitudinal assessment.

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By utilizing weightbearing scans, the foot can be assumed to have taken on its usual shape in response to load, and the relative positions of bones in this configuration may be used as a reference configuration to define initial static joint spacings, ligament lengths, and weightbearing kinematics.

11.6

Overview of biplane software

Specialized software is required to take the raw biplanar fluoroscopic image sequences, process them into a usable form, reconstruct the laboratory camera geometries, and to finally track objects in the stereo images to infer their 3D poses [61] (Fig. 11.3). The first steps in the software pipeline address the intensity inhomogeneities and spatial distortions imparted on the fluoroscopy data by the image intensifier and video lens. The intrinsic and extrinsic camera parameters that describe the biplane imaging chains are derived from a multi-frame dynamic calibration procedure that utilizes bundle adjustment to optimize for minimal reprojection errors [62]. Having performed this preprocessing, the virtual biplane software can then be used to track objects in 3D lab space. Marker-based tracking software detects the 2D centroids of X-ray attenuating fiducial marker spheres (beads) on the image pairs. Essentially a modern digitized form of the pioneering radiostereometric analyses, marker-based methods offer the highest degree of kinematic accuracy and have been used to quantify the motion of bones, implants, and soft tissues, however the invasive procedure of rigidly embedding markers into subjects has restricted in vivo use to special cases [28 31]. Model-based tracking software optimally aligns simulated radiographic projections of a volumetric bone model derived from a statistical model or via subject-specific imaging like CT or MRI. After an initial pose guess is provided in the virtually reconstructed biplane system, the volumetric bone models are perturbed in six degrees of freedom iteratively in an optimization routine, and simulated X-ray images, called digitally reconstructed radiographs (DRRs), are projected onto the rectified fluoroscopic image pairs (Fig. 11.4). The cost function of this optimization is driven by image similarity

FIGURE 11.3 The first step of model-based bone tracking involves rectifying the errors imparted on the fluoroscopy images by the image intensifier system and segmenting the bones to be tracked from computed tomography volumes. The biplane system is virtually recreated from a camera calibration procedure and the volumetric bone data manipulated in the reconstructed capture volume while ray-tracing algorithms generate simulated X-ray images called digitally reconstructed radiographs (DRR). The pose of the bone is reconstructed when the simulated X-ray projections maximally match the stereo fluoroscopy data.

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FIGURE 11.4 Model-based tracking can resolve the three-dimensional positions and orientations of the foot and ankle bones during human gait, including the phases of (A) initial contact, (B) midstance, and (C) terminal stance. The hindfoot and first ray motions are shown reconstructed (left) following the iterative model-based tracking alignment with the stereo fluoroscopy images (right).

metrics typically based on image intensity values or edges that correspond to osseous anatomy. Extensive user intervention is required during the model-based tracking process, as solutions may fall into local optima or be confounded by the presence of implants, lab hardware, sensors, shoes, and image noise and scatter in the fluoroscope sequences. While model-based tracking is more computationally expensive and slightly less accurate kinematically compared to markerbased tracking, it avoids invasive procedures of marker injection or bone pin fixation on test subjects. These restrictions are greatly reduced in cadaveric testing, and marker-based tracking has been used to validate the precision and accuracy of model-based tracking algorithms [20]. CT scans offer the best contrast between bony and non-bony tissues for segmentation. Segmentations of the foot and ankle bones from CT scans are most accurately performed manually or semi-manually, exploiting the knowledgebase of experienced technicians familiar with the bones of interest. However, manual segmentations are time consuming and can form processing bottlenecks in studies with large cohorts or many bones to track. Researchers have made efforts to automate the segmentation process with statistical shape models and machine learning methods for initializing new

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segmentations based on previously segmented training data [63]. These automated tools at the very least provide good initial segmentations that can be manually refined by technicians. Acceptable levels of CT bone segmentation errors have not been rigorously established for model-based tracking, particularly in the context of DRR similarity to the underlying fluoroscope images, anatomical coordinate systems based on the CT bone models, or arthrokinematic measures like joint spacing.

11.7

Clinical biplane foot and ankle studies

Biplane fluoroscopy has been employed to explore all types of clinical questions. As the kinds of analyses and variables from biplane systems grow, research thrusts can expand into clinical cohorts that have been underserved or seen limited research due to technological limitations. As a means of organizing previous work, studies will be grouped by the hardware employed.

11.7.1 Biplane systems consisting of two C-arms The research group at Massachusetts General Hospital (MGH) was one of the first teams to use biplane fluoroscopy to quantify foot bone motion (Fig. 11.5A). Their initial study reported normal ankle joint kinematics for five subjects in three static positions (heel strike, midstance, and toe off) [17]. Fluoroscopy data were collected on two OEC 9800 ESP General Electric (GE) C-arms and volumetric bone scans were collected on a 1.5 T GE MRI system. Bone position was manually manipulated in custom software until the projected cortical outlines matched the underlying fluoroscope images. The same methodology was extended to estimate ankle cartilage contact area [64] and cartilage strain [65]. Once the bones of interest from the MRI data were aligned, the overlap of the segmented cartilage models defined the articular contact area; average contact area was reported for nine subjects [64]. The depth of the cartilage overlap was used to calculate compressive strain, which was determined for six subjects [65]. This system was then extended to estimate the lengths of the anterior talofibular and calcaneofibular ligaments as the shortest distances between the centroids of the digitized attachment areas [66]. Finally, the same system was used to demonstrate aberrant ankle joint kinematics in six subjects with unilateral post-traumatic ankle arthritis [67]. A second group at Duke University used a system similar to the MGH setup consisting of two C-arms (Philips Pulsera) and an MRI (Siemens Trio 3 T) to study foot bone kinematics (Fig. 11.5B). They found increased translation and rotation in anterior talofibular ligament-deficient ankles compared to contralateral normal ankles as subjects stepped onto a level surface [18]. A similar methodology was employed for seven patients with unilateral lateral chronic ankle instability (CAI) [76]. Cartilage strain was calculated based upon the overlap of talar and tibial cartilage. CAI ankles were found to have increased strain compared to the contralateral control limb. Finally, seven patients underwent surgical repair for lateral ankle instability [68]. Post-surgical ankles exhibited significantly less motion while stepping onto a force plate; there was no difference between the post-surgical ankles and the contralateral normal ankles. A third research team, based at Fudan University in Shanghai, used dual fluoroscopes to study ankle motion, but they employed CT scans (GE Lightspeed) rather than MRI scans for the volumetric bone models [69]. Eleven subjects with CAI were enrolled and imaged with two BV Pulsera (Philips) C-arms (Fig. 11.5C) at 30 Hz while walking with and without an ankle brace along a flat floor and a 15-degree inversion platform. The silhouettes of the 3D bone models were semi-automatically matched to the fluoroscopy data at seven key poses during stance. Subtalar joint motion was more inverted when walking barefoot on the inversion platform; this motion decreased when the brace was worn. The same group used this technology to study ankle kinematics during five key poses of stair descent for ten controls, ten subjects with functional ankle instability (FAI), and ten lateral ankle sprain copers [77]. In general, the FAI subjects had excessive tibiotalar inversion and subtalar joint hypermobility, while the copers had a stable subtalar joint with tibiotalar inversion only at foot strike. Finally, a group from Western University in London, Ontario also employed a similar system consisting of two 23 cm fluoroscopes (SIREMOBIL Compact-L mobile C-arms, Siemens, Fig. 11.5D). Model-based bone tracking was conducted by generating 3D bone models from an a priori CT scan. The bone models were matched to the osseous edges in the X-ray data by manual manipulation using custom-developed software. This system was used to quantify foot shape (i.e., arch height defined by the calcaneus, navicular, and first metatarsal) across foot type both statically and dynamically for cavus, normal, and planus feet [70]. Next, the team applied the same technique to explore the effect of hard and soft custom foot orthoses using the same foot types and output parameters [78]. As compared to the barefoot and shod conditions, the arch shape was significantly different for the custom orthotic conditions, but not the OEM device.

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FIGURE 11.5 (A) Schematic of two OEC 9800 ESP General Electric C-arms at Massachusetts General Hospital [17]; (B) Two Philips Pulsera C-arms at Duke University [68]; (C) Schematic of two Philips BV Pulsera C-arms at Fudan University in Shanghai [69]; (D) Schematic of two Siemens SIREMOBIL Compact-L mobile C-arms at Western University in London, Ontario [70]; (E) Two modified Philips BV Pulsera C-arms at the Steadman Philippon Research Institute [71]; (F) Schematic of two modified Philips BV Pulsera C-arms at the VA Puget Sound; (G) Two modified GE OEC 9000 fluoroscopy systems at Marquette University [72]; (H) Two EMD Technologies pulsed X-ray generators (not shown) and two Shimadzu image intensifiers at Henry Ford Health Systems [22]; (I) Two Varian X-ray emitters and two Dunlee image intensifiers at the University of Utah [73]; (J) Schematic of a Gemss Medical biplane system at Chung-Ang University in Seoul, Republic of Korea [24]; (K) Schematic of two EMD X-ray generators and two Shimadzu image intensifiers at Brown University [74]; (L) Two EMD high-frequency cardiac cineradiographic generators and two 40 cm Thales image intensifiers at the University of Pittsburgh [75].

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11.7.2 Biplane systems consisting of two modified C-arms The systems from MGH, Duke, Shanghai, and London, ON were all essentially OEM C-arms with stock software that were positioned so that a subject could load their foot while being imaged. As such, there are some inherent hardware limitations (e.g., the physical imposition of the “C” support) that precluded uninhibited gait and software limitations (e.g., the sampling rate) that prevented blur-free imaging. To address these specific issues, groups based at the Steadman Philippon Research Institute, VA Puget Sound, and Marquette University developed systems that removed the image intensifiers and X-ray sources from the C-arms and mounted them on custom support stands. Further, the stock cameras were replaced by high-speed video cameras. Steadman Philippon tracked the motion of the calcaneus and tibia for six subjects walking barefoot and shod [19], demonstrating that subjects have different hindfoot motion while wearing shoes. They used BV Pulsera (Philips) C-arms with 30 cm image intensifiers that had been modified by removing the “C,” mounting on custom stands (Fig. 11.5E), and adding high-speed digital cameras (Phantom V5.1, Vision Research). Model-based RSA software was used to align 3D bone models with bony contours from the fluoroscopes. A second group with hardware similar to Steadman Philippon was the VA Puget Sound; they also employed modified BV Pulsera (Philips) C-arms with Phantom V5.1 high-speed cameras and custom support stands (Fig. 11.5F). Their initial work consisted of both a marker-based and a model-based validation for foot bone tracking [20,61]. Model-based bone tracking software was written in-house and used DRRs to match 3D bone positions with the 2D fluoroscopy data. A third research team from Marquette University has conducted several biplane studies with two modified GE OEC 9000 fluoroscopy systems with high-speed video cameras attached (Fig. 11.5G). By tracking two points per bone on the calcaneus and talus in the fluoroscopy data, and synchronizing a Vicon retro-reflective motion analysis system, kinematics of the tibiotalar and subtalar joint were determined for 13 subjects [21]. This group also reported the design, calibration, evaluation, and limitations of a single-plane version of their system [79], and explored the effect of crossscatter contamination in biplane fluoroscopy [72]. Next, the system was employed to quantify the effect of short and tall controlled ankle movement (CAM) boots on the talocrural and subtalar joints [80]. It was found that each CAM boot design limited motion at both joints. The group validated the bias and precision errors of their model-based tracking algorithm by tracking a cadaver foot embedded with fiducial beads through the system [81]. Finally, the system was used to compare barefoot to shod kinematics in 13 healthy subjects [82], demonstrating that biplane fluoroscopy could quantify bone motion while subjects wore shoes.

11.7.3 Biplane systems consisting of independent X-ray sources and image intensifiers A research team based at Henry Ford Health Systems used a biplane system consisting of 100 kW pulsed X-ray generators (EMD Technologies CPX 3100CV) and two 30 cm Shimadzu image intensifiers (AI5765 HVP, Fig. 11.5H) and custom DRR-based model tracking software [15] to study foot bone motion. The first study compared barefoot running to both an ultra flexible training shoe and motion control shoe, finding subtle differences in the tibiotalar and subtalar joints between the various conditions for 12 normal subjects during over ground running [22]. Next, the navicular drop height for the same three shoes was studied for 12 subjects during over ground running; there was no change in navicular drop magnitude but there was a slower navicular drop rate for the motion control shoes [83]. The University of Utah group has conducted several foot and ankle biplane studies, with the initial hardware setup consisting of two Varian X-ray emitters (Housing B-100/Tube A-142) and two 30 cm Dunlee image intensifiers (T12964-P/S), each mounted to a dedicated base around a treadmill. After an initial marker-based validation [23] of their bone tracking software [15], they used their biplane system to explore the potential for using retro-reflective skin markers for tracking tibiotalar and subtalar joint motion [84]. This system also elucidated the complex six degree-offreedom relationship between the tibiotalar and subtalar joints on ten normal subjects [85], demonstrating that the tibiotalar joint was responsible for sagittal plane motion, while the subtalar joint facilitated coronal and transverse plane motion, with some sagittal plane motion as well. Expanding their work to study four patients with CAI, they preliminarily demonstrated the aberrant kinematics of this pathologic population [86]. Reconfiguring their system around a force plate instead of a treadmill (Fig. 11.5I), the Utah group next described a methodology for quantifying tibiotalar kinematics after a total ankle replacement by leveraging computer-aided design models [87]. The reconfigured hardware system was then employed to retrospectively study ten subjects with an ankle arthrodesis [73]. By comparing to the unaffected contralateral limb, it was shown that subtalar joint plantarflexion is the primary compensatory mechanism after ankle arthrodesis. Finally, the biplane system was used to evaluate a multisegment foot model [88], demonstrating that kinematics derived from motion capture are similar to biplane-based motion data for the ankle and midfoot.

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Researchers based primarily at Chung-Ang University in Seoul, Republic of Korea, used a biplane system (KMC1400ST, Gemss Medical, Fig. 11.5J) and edge-based 3D/2D registration to track the tibia, talus, and calcaneus for ten healthy subjects [24]. They reported normal motion of the tibiotalar and subtalar joints. This system was next employed to study the relative movement of the articular surfaces of the tibiotalar and subtalar joints for 18 normal subjects [89]. Both joints demonstrated significant translational and rotational motion in early stance phase, but only the subtalar joint had significant rotational movement in late stance. The midtarsal joint locking mechanism was then explored on the same 18 subjects [90]. The data demonstrated that the midtarsal joint had continuous motion toward the extreme poses in terminal stance, failing to indicate any reduced motion or locking. Finally, this group used their biplane system to compare the hindfoot kinematics of five flatfoot subjects to the 18 normal subjects previously studied [91]. The flatfoot subjects had larger peak motions and ranges of motion at several joints, demonstrating abnormal kinematics during midstance and terminal stance. A collaborative research effort involving investigators from Australia, Canada, and the United States has conducted biplane foot studies with hardware consisting of two X-ray generators (EMD Technologies CPX 3100CV) and two image intensifiers (Shimadzu Medical Systems, model AI5765HVP, Fig. 11.5K), optically coupled to synchronized high-speed video cameras (Phantom IV, Vision Research). DRRs derived from CT scans were projected on the X-ray images and manually manipulated for each bone and participant using an open-source software package [92]. Their first study compared biplane videography to optical motion capture [74], and demonstrated that comparable solutions could be obtained in the sagittal plane, but out-of-plane motion should be carefully considered. Next, the team explored the intra- and inter-rater reliabilities of the system [93], and found root-mean-square errors from 2 to 3 mm and 2 to 3 degrees. Finally, the team used the biplane system to validate a constrained multi-segment foot model based on OpenSim [94]. The model was contrasted to a conventional multi-segment foot model and found to be more accurate. Building on their long track record in biplane fluoroscopy, the research team at the University of Pittsburgh has applied their expertise to the foot and ankle. Using a system that consists of two 150 kV constant potential highfrequency cardiac cineradiographic generators (EMD CPX-3100CV, EMD Technologies) and two high-speed digital cameras (4 megapixel Phantom V10; Vision Research, Wayne, NJ) coupled to 40 cm Thales image intensifiers (Fig. 11.5L). Two subjects had metal beads implanted in their tibia, fibula, talus, and calcaneus, and then a bead-based validation of the bone-based tracking was conducted by automatically manipulating DRRs to maximize image correlation [95]. Errors were less than 1 mm in translation and less than 2 degrees of rotation. Next, this group evaluated four techniques for generating hindfoot bone coordinate systems [75]. Rotations as defined by these systems were compared to the functional axes (i.e., helical axes of rotation) determined from biplane radiography. Finally, this team explored kinematic differences at the tibiotalar and subtalar joints due to sex and dominant versus non-dominant limb in 20 healthy adults [96]. Additionally, several other research teams have conducted biplane foot and ankle studies. A group from Keio University has used a custom biplane system with X-ray sources and FPDs to study cadaveric feet in static loading conditions [97 99]. A Taiwanese research team has developed a refined two-step bone tracking technique that they validated on a single cadaver foot [100]. Another group has used two BV Pulsera C-arms to quantify the material properties of the heel pad [101]. There are also at least two published literature reviews on foot and ankle biplane studies [100,102].

11.8

Future applications and directions

The typical goal of a biplane study is to generate six degree-of-freedom kinematic descriptions of joint motion (Fig. 11.6A). These are reported using anatomically based coordinate systems embedded in the bones of interest [59,75]. For many researchers, the precise kinematic data possible with biplane systems are enough on their own to provide insights into subtle differences or changes over time in joints in response to clinical interventions or degradations. However, joint kinematics are only a subset of the possible variables that may be generated based on biplane data. Having accurately tracked the positions of two bones that comprise a joint, arthrokinematic measures of articular interactions (Fig. 11.6B) such as the centroid of contact, relative shear velocity, and dynamic measurements of joint space and congruity are feasible [23]. By coupling subject-specific definitions of ligamentous insertions and origins from MRI, CT scans, or statistical atlases with rudimentary models of the ligament geometry, tracking skeletal motion affords estimates of dynamic ligament deformation patterns (Fig. 11.6C) and offers another avenue of investigation in the realm of soft tissues [103]. Biplane systems have also been applied for tracking the deformation of the heel pad and

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FIGURE 11.6 Some of the variables obtainable from biplane fluoroscopy. (A) Six degree-of-freedom kinematic descriptions of joint motion; (B) Arthrokinematic measures of articular interactions such as centroid of contact or dynamic measurement of joint space; (C) Estimates of dynamic ligament deformation patterns.

plantar tissues in vivo [103]. Moreover, as both biplane hardware and software mature, researchers and clinicians can address fundamental and foundational questions on foot function.

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Model-based tracking of the bones of the foot: a biplane fluoroscopy validation study. Comput Biol Med 2018;92:118 27. [62] Lichti DD, Sharma GB, Kuntze G, Mund B, Beveridge JE, Ronsky JL. Rigorous geometric self-calibrating bundle adjustment for a dual fluoroscopic imaging system. IEEE Trans Med Imaging 2015;34:589 98. [63] Brehler M, Islam A, Vogelsang L, Yang D, Sehnert W, Shakoor D, et al. Coupled active shape models for automated segmentation and landmark localization in high-resolution CT of the foot and ankle. Proc SPIE Int Soc Opt Eng 2019;10953. [64] Wan L, de Asla RJ, Rubash HE, Li G. Determination of in-vivo articular cartilage contact areas of human talocrural joint under weightbearing conditions. Osteoarthr Cartil 2006;14(12):1294 301. [65] Wan L, de Asla RJ, Rubash HE, Li G. In vivo cartilage contact deformation of human ankle joints under full body weight. J Orthop Res 2008;26(8):1081 9. [66] de Asla RJ, Kozanek M, Wan L, Rubash HE, Li G. Function of anterior talofibular and calcaneofibular ligaments during in-vivo motion of the ankle joint complex. J Orthop Surg Res 2009;4:7. [67] Kozanek M, Rubash HE, Li G, de Asla RJ. Effect of post-traumatic tibiotalar osteoarthritis on kinematics of the ankle joint complex. Foot Ankle Int 2009;30(8):734 40. [68] Wainright WB, Spritzer CE, Lee JY, Easley ME, DeOrio JK, Nunley JA, et al. The effect of modified Brostrom-Gould repair for lateral ankle instability on in vivo tibiotalar kinematics. Am J Sports Med 2012;40(9):2099 104. [69] Cao S, Wang C, Zhang G, Ma X, Wang X, Huang J, et al. Effects of an ankle brace on the in vivo kinematics of patients with chronic ankle instability during walking on an inversion platform. Gait Posture 2019;72:228 33. [70] Balsdon ME, Bushey KM, Dombroski CE, LeBel ME, Jenkyn TR. Medial longitudinal arch angle presents significant differences between foot types: a biplane fluoroscopy study. J Biomech Eng 2016;138(10).

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[71] Giphart JE, Zirker CA, Myers CA, Pennington WW, LaPrade RF. Accuracy of a contour-based biplane fluoroscopy technique for tracking knee joint kinematics of different speeds. J Biomech 2012;45(16):2935 8. [72] Cross JA, McHenry B, Schmidt TG. Quantifying cross-scatter contamination in biplane fluoroscopy motion analysis systems. J Med Imaging (Bellingham) 2015;2(4):043503. [73] Lenz AL, Nichols JA, Roach KE, Foreman KB, Barg A, Saltzman CL, et al. Compensatory motion of the subtalar joint following tibiotalar arthrodesis: an in vivo dual-fluoroscopy imaging study. J Bone Joint Surg Am 2020;102(7):600 8. [74] Kessler SE, Rainbow MJ, Lichtwark GA, Cresswell AG, D’Andrea SE, Konow N, et al. A Direct comparison of biplanar videoradiography and optical motion capture for foot and ankle kinematics. Front Bioeng Biotechnol 2019;7:199. [75] Brown JA, Gale T, Anderst W. An automated method for defining anatomic coordinate systems in the hindfoot. J Biomech 2020;109:109951. [76] Bischof JE, Spritzer CE, Caputo AM, Easley ME, DeOrio JK, Nunley JA, et al. In vivo cartilage contact strains in patients with lateral ankle instability. J Biomech 2010;43(13):2561 6. [77] Cao S, Wang C, Ma X, Wang X, Huang J, Zhang C, et al. In vivo kinematics of functional ankle instability patients and lateral ankle sprain copers during stair descent. J Orthop Res 2019;37(8):1860 7. [78] Balsdon M, Dombroski C, Bushey K, Jenkyn TR. Hard, soft and off-the-shelf foot orthoses and their effect on the angle of the medial longitudinal arch: a biplane fluoroscopy study. Prosthet Orthot Int 2019;43(3):331 8. [79] McHenry BD, Exten E, Long JT, Harris GF. Sagittal fluoroscopy for the assessment of hindfoot kinematics. J Biomech Eng 2016;138 (3):4032445. [80] McHenry BD, Exten EL, Cross JA, Kruger KM, Law B, Fritz JM, et al. Sagittal subtalar and talocrural joint assessment during ambulation with controlled ankle movement (CAM) boots. Foot Ankle Int 2017;38(11):1260 6. [81] Cross JA, McHenry BD, Molthen R, Exten E, Schmidt TG, Harris GF. Biplane fluoroscopy for hindfoot motion analysis during gait: a modelbased evaluation. Med Eng Phys 2017;. [82] McHenry BD, Kruger KM, Exten EL, Tarima S, Harris GF. Sagittal subtalar and talocrural joint assessment between barefoot and shod walking: a fluoroscopic study. Gait Posture 2019;72:57 61. [83] Hoffman SE, Peltz CD, Haladik JA, Divine G, Nurse MA, Bey MJ. Dynamic in-vivo assessment of navicular drop while running in barefoot, minimalist, and motion control footwear conditions. Gait Posture 2015;41(3):825 9. [84] Nichols JA, Roach KE, Fiorentino NM, Anderson AE. Predicting tibiotalar and subtalar joint angles from skin-marker data with dualfluoroscopy as a reference standard. Gait Posture 2016;49:136 43. [85] Roach KE, Wang B, Kapron AL, Fiorentino NM, Saltzman CL, Bo Foreman K, et al. In vivo kinematics of the tibiotalar and subtalar joints in asymptomatic subjects: a high-speed dual fluoroscopy study. J Biomech Eng 2016;138(9). [86] Roach KE, Foreman KB, Barg A, Saltzman CL, Anderson AE. Application of high-speed dual fluoroscopy to study in vivo tibiotalar and subtalar kinematics in patients with chronic ankle instability and asymptomatic control subjects during dynamic activities. Foot Ankle Int 2017;38 (11):1236 48. [87] Blair DJ, Barg A, Foreman KB, Anderson AE, Lenz AL. Methodology for measurement of in vivo tibiotalar kinematics after total ankle replacement using dual fluoroscopy. Front Bioeng Biotechnol 2020;8:375. [88] Roach KE, Foreman KB, MacWilliams BA, Karpos K, Nichols J, Anderson AE. The modified Shriners Hospitals for Children Greenville (mSHCG) multi-segment foot model provides clinically acceptable measurements of ankle and midfoot angles: a dual fluoroscopy study. Gait Posture 2021;85:258 65. [89] Phan CB, Nguyen DP, Lee KM, Koo S. Relative movement on the articular surfaces of the tibiotalar and subtalar joints during walking. Bone Joint Res 2018;7(8):501 7. [90] Phan CB, Shin G, Lee KM, Koo S. Skeletal kinematics of the midtarsal joint during walking: midtarsal joint locking revisited. J Biomech 2019;95:109287. [91] Phan CB, Lee KM, Kwon SS, Koo S. Kinematic instability in the joints of flatfoot subjects during walking: a biplanar fluoroscopic study. J Biomech 2021;127:110681. [92] Brainerd EL, Baier DB, Gatesy SM, Hedrick TL, Metzger KA, Gilbert SL, et al. X-ray reconstruction of moving morphology (XROMM): precision, accuracy and applications in comparative biomechanics research. J Exp Zool A Ecol Genet Physiol 2010;313(5):262 79. [93] Maharaj JN, Kessler S, Rainbow MJ, D’Andrea SE, Konow N, Kelly LA, et al. The reliability of foot and ankle bone and joint kinematics measured with biplanar videoradiography and manual scientific rotoscoping. Front Bioeng Biotechnol 2020;8:106. [94] Maharaj JN, Rainbow MJ, Cresswell AG, Kessler S, Konow N, Gehring D, et al. Modelling the complexity of the foot and ankle during human locomotion: the development and validation of a multi-segment foot model using biplanar videoradiography. Comput Methods Biomech Biomed Engin 2021;1 12. [95] Pitcairn S, Kromka J, Hogan M, Anderst W. Validation and application of dynamic biplane radiography to study in vivo ankle joint kinematics during high-demand activities. J Biomech 2020;103:109696. [96] Yang S, Canton SP, Hogan MV, Anderst W. Healthy ankle and hindfoot kinematics during gait: sex differences, asymmetry and coupled motion revealed through dynamic biplane radiography. J Biomech 2021;116:110220. [97] Ito K, Hosoda K, Shimizu M, Ikemoto S, Kume S, Nagura T, et al. Direct assessment of 3D foot bone kinematics using biplanar x-ray fluoroscopy and an automatic model registration method. J Foot Ankle Res 2015;8:21. [98] Ito K, Hosoda K, Shimizu M, Ikemoto S, Nagura T, Seki H, et al. Three-dimensional innate mobility of the human foot bones under axial loading using biplane x-ray fluoroscopy. R Soc Open Sci 2017;4(10):171086.

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[99] Negishi T, Nozaki S, Ito K, Seki H, Hosoda K, Nagura T, et al. Three-dimensional innate mobility of the human foot on coronally-wedged surfaces using a biplane x-ray fluoroscopy. Front Bioeng Biotechnol 2022;10:800572. [100] Lin CC, Li JD, Lu TW, Kuo MY, Kuo CC, Hsu HC. A model-based tracking method for measuring 3D dynamic joint motion using an alternating biplane x-ray imaging system. Med Phys 2018;. [101] Teng ZL, Yang XG, Geng X, Gu YJ, Huang R, Chen WM, et al. Effect of loading history on material properties of human heel pad: an in-vivo pilot investigation during gait. BMC Musculoskelet Disord 2022;23(1):254. [102] Canton S, Anderst W, Hogan MV. In vivo ankle kinematics revealed through biplane radiography: current concepts, recent literature, and future directions. Curr Rev Musculoskelet Med 2020;13(1):77 85. [103] Kroupa N, Pierrat B, Han WS, Grange S, Bergandi F, Molimard J. Bone position and ligament deformations of the foot from CT images to quantify the influence of footwear in ex vivo feet. Front Bioeng Biotechnol 2020;8:560.

Chapter 12

Plantar Pressure and Ground Reaction Forces Dieter Rosenbaum1 and Scott Telfer2,3,4 1

Institute of Experimental Musculoskeletal Medicine, Movement Analysis Laboratory, University Hospital Muenster, Muenster, Germany, 2Department

of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States, 3Department of Mechanical Engineering, University of Washington, Seattle, WA, United States, 4RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States

Abstract Assessing the forces between the foot and the ground is a key part of many biomechanical analyses of locomotion. Studying these loading patterns can help us to understand how the forces are transferred to the tissues of the body. These forces are essential for maintaining the health of tissues, but they can lead to damage and injury if they are extremely high. There are several different tools available to capture these types of data, and this chapter will focus on pressure measurement systems and force platforms. We will cover topics including the development of these tools, the analysis of the data generated, and give examples of the use of these types of data in the context of diabetic foot disease, foot type, and sports.

12.1 Introduction: clinical relevance of force and pressure measurements in foot and ankle biomechanics During bipedal human locomotion, the foot is the terminal link in the kinetic chain and as such, it is in contact with the ground and exchanges forces according to Newton’s third law “actio 5 reactio.” This entails that the forces being exerted against the ground during landing and propulsion are causing reactive forces that are simultaneously transferred into the body with equal amplitude and opposite direction. The external forces that are acting on the human body are transferred to the adjoining anatomical structures and to the proximal segments, thus contributing to the internal loads in all structures including the bones, joints, cartilage, muscles, tendons, and ligaments. Usually, these loads can be sustained by the tissues as long as they fall within a certain range that is determined by the magnitude, frequency, and duration of the loading events. Furthermore, a certain amount of loading is usually necessary to maintain the tissues and structures as they react to external stress by remodeling processes [1]. If the loads stay below a certain threshold, tissues are reduced in the body’s effort to adjust to the reduced external demands. Above-threshold loads, on the other hand, may cause positive adaptations to some degree but may also eventually lead to overload injuries or problems when the adaptive or remodeling processes are too slow or insufficient for withstanding the increased loading. While these processes are generally well described and accepted, the actual loading thresholds for “just enough” or “not too much” are not clearly defined as they also appear to be individually determined to a certain degree. Therefore it has always been of interest to investigate the external forces acting on the human body during the whole range of locomotor activities. This research has helped to advance the understanding of physiological and pathological loading in healthy subjects as well as in patients suffering from structural and/or functional deficits. For bipedal human locomotion, the two main areas of interest are the following: (1) plantar pressure distribution measurements (also termed pedography, pedobarography, or, less commonly, baropodometry) and (2) ground reaction Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00037-8 © 2023 Elsevier Inc. All rights reserved.

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force measurements. In this chapter, both research fields will be explained and illustrated with respect to measurement technology and analytical options.

12.2

Background: force versus pressure

There are two domains that ought to be distinguished in this field of research: ground reaction forces and plantar pressure measurements. Ground reaction forces describe the three-dimensional exchange of forces between the body and the ground, usually summarized as a single vector of changing magnitude and direction for each foot during the ground contact phase. Pressure distribution measurements, on the other hand, describe the distribution of these forces across the changing contact area during the roll-over process. The additional spatial information usually requires the shear force components to be “sacrificed” and only the vertical component of the ground reaction force is measured for further analysis. Paul W. Brand phrased the statement that Pressure is the critical quantity that determines the harm done by the force [2]. This distinguishes the essential relevance of these two physical quantities. This should help to understand that forces are driving the movements of the body and its segments during locomotion while the contact surfaces between the body and the environment experience the local effect of the forces as contact pressures that they have to sustain. In physics terms, the pressure P is defined and calculated as the force F divided by the contact area A. Its SI-derived unit is Pa (Pascal; 1 Pa 5 1 N/m2). The values for plantar pressure measurements are usually in the range of several hundred kPa up to slightly over 1 MPa (i.e., 1000 to 1 million Pascal, with 10 kPa equaling 1 N/cm2).

12.2.1 History The forces exerted on the human body during motion have been of interest to movement analysts from the very beginning. As early as 350 BC, Aristotle (384 322 BC) wrote his work De motu animalium (On the Motion of Animals) and described his observations of gait and movement in animals and humans. Since then, the scientific interest in the interaction between the human body and the ground has never faded. With the advent of modern electronics, researchers have been able to measure these forces dynamically and study their effects during a wide range of activities from normal locomotion to movement pathologies or impairments as well as athletic performance. The first systematic investigations of the load distribution during human walking were reported in the 19th century with recordings of the foot-to-ground contact in plaster of Paris and clay [3]. However, similar to the footprints left in the soft sand on a beach, this results in the information being summated over the whole ground contact and does not distinguish the distinct phases of the roll-over process from the initial heel contact to the push-off from the toes. Slightly more detailed information became available in the early 20th century when walking over steel balls left indentations in a lead plate [4] or foot-loaded rubber pyramids left ink prints on paper [5] with the size of the respective marks being directly related to the applied load. The first dynamic recordings were based on a similar pyramidal approach; however, the deformation was filmed (i.e. dynamically recorded) through a glass plate so that the change of loading could be analyzed with respect to the different phases of the ground contact [6]. For a routine application of plantar pressure measurements, a steadily increasing interest can be observed since the first commercially available systems were marketed in the 1980s. The first of these was based on a measurement principle developed by Nicol and Hennig with a capacitance device arranged in 16 rows and 16 columns of conducting strips [7]. These strips were separated by an elastic dielectric with each intersection forming one of 256 capacitors. By means of a multiplexing scheme, these sensors could be sampled at a frequency of 100 Hz. Based on this construction principle, the first EMED system was developed and marketed by Novel Inc. (Munich, Germany). Pressure measurement systems can be loosely divided into three types: platforms, mats, and in-shoe devices (Fig. 12.1). Platform systems are rigid plates containing an array of sensors that are semi-portable and are placed on the ground for people to walk over. The size of these systems determines how many steps can be collected in a single trial, often it is limited to just one. Mat systems consist of a thin, flexible sheet of sensors that can come in a wide range of sizes and can conform to some extent to nonplanar surfaces. In-shoe systems are similar to pressure mats but are designed specifically to fit in shoes and measure the full plantar surface. They generally are available in a range of standard shoe sizes or can be cut to fit different footwear. More portable than platform systems, in-shoe systems allow recording of multiple steps a person takes. Some more recent in-shoe systems allow wireless data logging, meaning extensive real-world activity can be collected. Over time, more measurement platforms, mats, and in-shoe systems have become commercially available, varying in sensor-type, material, spatial and temporal resolution, and overall size. Therefore a critical appraisal of the technical

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FIGURE 12.1 Platform (A) and in-shoe plantar pressure measurement systems (B).

specifications (discussed briefly later in this chapter) and the inherent advantages and disadvantages is warranted nowadays before the application of any of these systems to answer a research or clinical question. The differences between the technical properties and sensor characteristics also makes it difficult to compare research results obtained from different measurement systems [8]. Force platforms (or plates) are in some ways simpler in design, a result of only being intended to measure the reaction force when someone walks over them. Early platforms were based on pneumatic sensors, and then advanced to use strain gage and piezoelectric transducers. Early platforms only measured the vertical component of the ground reaction force, and it was not until the 1940s that platforms were developed to measure the full six components of the ground reaction force along with the location of the center of pressure. The first commercially available modern force platform was produced by Advanced Mechanical Technology Inc. in 1976 and installed at Boston Children’s Hospital [9]. For more information about the early development of force platforms, please refer to the review article by Baker [10]. Most modern force platforms for gait and movement analysis are in the form of a highly rigid rectangular plate with force transducers in each corner. The unit is embedded securely in the floor of the laboratory and isolated from the surrounding structures to avoid crosstalk (Fig. 12.2). When a subject walks across the plate, the outputs from the transducers are combined to provide an overall 3-dimensional ground reaction force vector with its center of pressure coordinates relative to the plate, along with the resulting free moments. Force platforms for collecting ground reaction forces are almost exclusively platform systems, although some force-measuring shoes have been developed [11]. Similar to plantar pressure measurement systems, care should be taken when assessing the technical specifications of different force platforms to find the appropriate system for different applications.

12.2.2 Development of measurement technologies Forces and pressures can be measured with dedicated sensors. While early sensors operated only mechanically (e.g., the foot leaving imprints in deformable material like plaster of Paris, clay, or sand; rubber mats with pyramidal structures increasing their contact area while being pushed on a glass plate; springs that deformed under load application), modern systems generally use electrical sensors. These sensors are usually based on resistive, inductive, capacitive, or piezoelectric technology. All of these sensor technologies have in common the characteristic that when an external load (force or pressure) is exerted on the sensor this changes the sensor’s electrical properties, that is, their resistance, inductance, capacitance, or charge distribution. This change can be transduced and amplified into an appropriate change in voltage which can then be displayed and recorded with dedicated monitors or software applications (after analog-to-digital conversion). For most sensor materials, a certain degree of physical deformation is necessary to achieve this effect. The exception is piezoelectric sensors, which use materials with a high modulus of elasticity and show almost no deflection under load. The applied loads affect the charge distribution in the

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FIGURE 12.2 An example of a modern motion analysis laboratory with an array of force platforms embedded in the floor.

crystal lattice by separating the positive and negative charges within so that the charge difference on the surface of the crystal can be amplified and converted into a voltage. For a more detailed discussion of advantages and disadvantages of specific sensors the reader is referred to the available technical review papers [12 14]. Before using any single sensor or array of multiple sensors for measurement applications, a calibration process is necessary to determine the relationship between the load input and the sensor output across the total range of expected values. In the case of pressure measurement systems, this process should be performed for each single sensor as the sensor characteristics might not be uniform across a sensor matrix. A single sensor calibration approach will result in an input-output curve that can be applied before data analysis so that meaningful and reliable force or pressure values can be extracted. This approach is preferable to calibrating across a whole sensor matrix by dividing a known force (as measured, e.g., by an underlying force plate) over the number of sensors of the matrix. While the overall force-load relationship might be more or less correct, the actual pressure values cannot be trusted due to the inherent sensor differences. With respect to sensor characteristics, there are several potentially critical technical features that should be considered when selecting the appropriate measurement technology for any given application. Perhaps the most important of these in the context of biomechanics are linearity, hysteresis, and temperature drift. G

G

G

Linearity describes the relationship between mechanical input, that is, force or pressure, and the sensor’s output signal. This should ideally be linearly related across the intended measurement range. However, some sensor materials might change their sensitivity with varying levels of load so that a nonlinear, more complex relationship might exist. This requires a deliberate calibration process taking the deviation from linearity into account. Hysteresis describes the difference in the sensor output between loading and unloading that may be due to the elasticity and responsiveness of the materials involved. The effect is usually frequency-dependent, that is, it is more pronounced in fast loading-unloading events. The effect is limited or can be neglected in stiff materials (e.g., piezoelectric crystals) and may be more pronounced in highly elastic materials (e.g., the dielectric in capacitive sensors). Temperature drift is a change in the sensor output with changing sensor temperatures and it has been reported especially in piezoelectric materials. For example, pressure measuring insoles will adapt to the body temperature when worn inside shoes for an extended period of time. Therefore if these insoles use sensors that have a pronounced temperature drift they should be given some time to warm up until a steady state has been reached before recording measurements.

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When these characteristics appear insufficiently controlled, they may have to be compensated for by appropriate signal conditioning and amplifier characteristics. Those sensors that sustain a more pronounced deformation under loading (primarily inductive sensors; to a lesser degree, capacitive sensors) might not recoil to the unloaded state quickly enough in highly dynamic or high-frequency loading situations such as jumping, thus causing a more pronounced hysteresis effect. Piezoceramic sensors are based on a charge separation in a stiff crystal lattice that deforms less meaning hysteresis effects are usually negligible. However, these sensors are more prone to altering their response characteristics under prolonged loading that may change the temperature of the piezo element thus causing a certain temperature drift. These examples illustrate that the user needs to be informed about the anticipated loading conditions of their application so that the most appropriate sensor technology can be selected and applied. For plantar pressure measurements specifically, the sensor itself should not affect the quantity that is being measured, that is, the pressure distribution. Therefore sensor materials should be thin enough not to provide a cushioning effect that might already change the acting pressure. Furthermore, it is desirable that the sensor covers and thus measures the total contact area in order not to miss potentially critical areas which might be at risk for overloading, especially in more pronounced foot deformities. The actual pressure values should be identified as accurately as possible and not be underestimated by under-sampling, that is, by using an insufficient spatial resolution of the sensors. This effect has been investigated by Davis et al. (1996) and was proven to be critical for sensor dimensions exceeding a size of 6.36 by 6.18 mm (mediolateral and anteroposterior directions) [15]. Further investigation was performed by Pataky who found that critical sensor widths could range from 1.7 to 17.4 mm depending on the measurement scenario [16]. Larger sensors might have the effect that high-pressure peaks with a small dimension would get averaged across the surrounding areas of lower pressures so that the actual peak pressure value would be underestimated. This effect becomes more pronounced with increasing sensor size as well as with more pronounced and localized pressure peaks and high-pressure gradients.

12.2.3 Visualization and analytical options Pressure distribution patterns under the foot are most often displayed in a color scale with the low-pressure zones being presented in “cold”, that is, blue colors changing toward “hot”, that is, red or pink colors for high-pressure zones (Fig. 12.3). The presentation can be based on pixels representing each single sensor or it can be extrapolated and smoothed between the single sensors so that anatomical structures become more easily recognizable. These color maps

FIGURE 12.3 A typical visualization of plantar pressure data, colorized to allow easy interpretation and identification of high-pressure areas.

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FIGURE 12.4 Example of masking scheme for plantar pressure data. Regions are hindfoot, midfoot, metatarsal heads 1 5, hallux and lesser toes. Center of pressure line is also included.

usually provide a valuable first impression about the pressure pattern indicating where the plantar surface is predominantly loaded during ground contact. While this helps to intuitively understand the loading characteristics, it allows only for a first and coarse analysis as it does not consider the actual pressure values. For a quantitative analysis, however, the actual values should be analyzed. Owing to the complexity of the task— given that a foot with a plantar surface area of around 100 cm2 would be in contact with approximately 400 sensors (for a given spatial resolution of 4 sensors per cm2)—and to allow comparisons of the data between and across individuals, a regional analysis of the whole footprint is usually applied. For this purpose, the foot is subdivided into pre-defined regions of interest, areas or so-called masks that are based on the foot anatomy (Fig. 12.4). The degree of subdivision, that is, the number of areas, depends on the research question, and as such there is not one generally accepted, or “best” approach, but one may choose from several different options. For proximal to distal, it usually makes sense to distinguish between hindfoot, midfoot, forefoot and toe regions and these regions can—for most feet—be defined manually

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or, preferably, by fixed percentages of the total foot length or by shape recognition methods. For medial to lateral, the hindfoot and midfoot may (or may not) be subdivided into a medial and lateral half. In the forefoot, the question is whether the five metatarsal heads can be separately analyzed since the plantar fat pad usually prevents five individual pressure peaks from being distinguished. Therefore it might be enough to define and distinguish a medial (i.e., 1st metatarsal), central (2nd and 3rd metatarsal) and lateral (4th and 5th metatarsal) forefoot region. For the toes, the hallux can be easily identified as a large, oval-shaped, well-loaded structure whereas the lesser toes, except for the often protruding second toe, contribute less to the final push-off from the ground. In the literature, masks ranging from 2 to 12 regions can be found. As stated before, the decision for a mask to be applied should be based on the degree of information that is required for answering the clinical question at hand. Furthermore, this decision also depends to some degree on the expected degree of deformity as some patients with severely affected feet might not contact the ground with the whole plantar surface. In these cases, the automatic masking procedures or algorithms might not be successfully applied. Another example would be when the heel contact is lacking in patients with shortened Achilles’ tendons. This might require manual masking to define “empty” masks for those regions that are not in contact with the ground. The next issue is the selection of parameters that can be extracted for the selected foot regions. The first and most straightforward parameter is usually the highest or peak pressure and/or the average pressure acting in a certain region of interest. Peak pressures report only the maximum value of the single sensor that experienced the highest load during the contact phase. Average pressures, on the other hand, determine the average across all sensors in the region of interest. From the authors’ point of view, the peak pressure is clinically more relevant because it determines whether the loading might be felt as uncomfortable or painful. Nevertheless, it might make sense to determine the total load in a region of interest so that regional forces, that is, the summation of the pressures multiplied by the contact area can be calculated. This allows us to determine how the vertical component of the ground reaction force during ground contact is distributed over the respective foot regions. Furthermore, the size of the foot regions, that is, contact areas and the duration of ground contact, that is, contact times, can be compared. For the overall foot loading, a valuable parameter is the regional force-time-integral which has been defined in different ways in the literature [17] but takes both the amplitude and the duration of the load application into account. These parameters can be used with their absolute values or can be normalized with respect to the total load, area, or time, to make them comparable even for feet or subjects with different anthropometrics. Regional forces can be expressed as a percentage of the total bodyweight, regional contact areas as percentages of the total foot contact area, regional contact times as percentages of the total ground contact time, and regional force-time-integrals as percentages of the total impulse. For the peak or average pressure values, a normalization does not appear warranted since pressures are not linearly related to the bodyweight and depend on further experimental conditions. It should be noted that foot loading is dependent on walking speed since a higher speed increases the gait dynamics and will generally lead to higher forces and pressures underneath the foot, even though this effect is not uniformly seen under all foot regions [18,19]. Therefore walking speed should be monitored during a recording session and preferably be comparable for example if patients are to be evaluated repeatedly, for example, before and after therapy. This should help to ensure that differences in foot loading characteristics are due to changes in clinical conditions and not just due to different gait dynamics. These analysis approaches have led to many insights into plantar pressure data, however, they tend to provide only a limited resolution analysis of the data, that is, the data for each region is summarized into a single, discrete value, and the use of anatomical masks means that the results can be sensitive to region boundary definitions [20]. More recently, researchers have begun to use tools based on statistical parametric mapping to allow a more global analysis of the full pressure dataset [21] (Fig. 12.5). Data from force plates can be analyzed as individual components (vertical, mediolateral, anteroposterior) (Fig. 12.6). Variables such as peak load or loading rate can be determined from these data, and specific tasks may have characteristic waveforms with discrete variables that can be compared between groups or individuals, for example, peaks and troughs in the “double hump pattern” typically seen in walking. As with plantar pressure data, force data may be normalized to bodyweight, and more advanced analysis techniques such as statistical parametric mapping [22] or principal component analysis [23] may be used to perform a fuller comparison between conditions. Force platforms have also been used to measure postural stability by tracking the excursion and velocity of the center of pressure during quiet standing [24]. However, only limited clinical or research-relevant information can be determined from the force plate data alone, and it is commonly used as an input for more complex biomechanical analyses, for example inverse dynamic studies that look at joint moments and musculoskeletal modeling simulations that study overall joint loading.

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FIGURE 12.5 Example results from statistical parametric analysis of plantar pressure data looking at changes in pressures between the ages of 4 7 years. (A) The raw statistical parametric map (SPM). (B) Inference images, depicting the portions of the feet which reached (or failed to reach) statistical significance. Phethean J, Pataky TC, Nester CJ, Findlow AH. A cross-sectional study of age-related changes in plantar pressure distribution between 4 and 7 years: a comparison of regional and pixel-level analyses. Gait Posture 2014;39:154 60.

FIGURE 12.6 Typical ground reaction force measurements from an individual walking over a force platform. AP, Anterior-posterior; ML, mediolateral; V, vertical.

12.3

Research applications and selected clinical examples

The continued interest in the application of pressure distribution measurements for research applications questions is illustrated in the number of publications with increasing contributions on a yearly basis (Fig. 12.7). A recently performed literature search on Scopus with the search terms “plantar pressure”, “foot pressure”, “pedograph*”, “pedobarograph*”, and “baropodo*” should reveal most publications in this field and resulted in over 5500 hits (as of January 2021). This indicates that the application of pressure distribution measurements nowadays is well established and widely used for answering a variety of clinical and research questions. The titles reveal the wide range of topics, with some more prominent.

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FIGURE 12.7 Number of publications per year related to plantar pressures.

With respect to relevance, the main topics deal with clinical issues such as diabetic foot problems or foot/toe deformities. On the other hand, sports-related topics are also widely investigated. Finally, shoe, footwear, and orthotics make up the third area of application that covers both, clinical as well as sports aspects. Some of these topics will be specifically addressed in the following paragraphs.

12.3.1 Diabetes One of the main topics, which has also driven some of the early technical developments of the first commercially available pressure measuring systems, is the functional assessment of patients with diabetes and their commonly encountered foot problems. Due to the neuropathy and angiopathy that many patients with diabetes develop over the course of their disease the pressures and forces the foot is exposed to are not reliably sensed by the peripheral neuro-receptors and therefore not registered by the central nervous system. Thus, repeated overloading may occur that will not be prevented by the insensitive patient since minor injuries may remain undetected in these individuals. Usually, the sensory feedback from nerve endings in the skin of these high-load areas would cause a centrally elicited offloading response due to pain or discomfort. However, this self-protecting mechanism is corrupted in patients with diabetes. Combined with changes in plantar tissue mechanics [25] that lead to higher peak pressures [8], these individuals are at a high risk of developing an ulcer. Therefore patients are advised to wear appropriate footwear to prevent the development of foot ulcerations. These most often occur in the areas of high local stress but also in less loaded areas like the tips of the toes. This surface stress can be measured reliably under the plantar surface with pressure distribution instrumentation, even with previous foot ulceration [26]. This can be achieved with two different approaches: 1. with a platform device for recording the dynamic loading during barefoot walking. 2. with an in-shoe device measuring the interaction between foot and shoe during shod walking. The latter approach can be used to determine whether a shoe is appropriate for the patient’s individual foot morphology and gait dynamics or whether certain shoe/insole modifications are advisable to optimize the pressure distribution by offloading areas at risk of overloading. This is especially effective with customized insoles that are based on the individual’s foot shape and that should be evaluated to ensure that they achieve the desired degree of offloading. Since peak pressure describes a potentially harmful effect for the local tissues underneath, this would be the main parameter of interest for assessing an individual patient’s risk. Therefore any countermeasures such as protective footwear and/or load-redistributing insoles should be evaluated with respect to their pressure-reducing effects in highly loaded foot regions (Fig. 12.8). In Germany, for example, the orthopedic shoemakers who provide such a service are required to document the offloading effect to be reimbursed from health insurers. Thus, they measure before and after insole insertion to verify that the local pressure has been reduced by a sufficient amount (with some ongoing discussion regarding what can be deemed “sufficient”). In general, the effectiveness of insoles for redistributing plantar pressure in diabetic patients has been well documented (e.g., Tsung et al. [27]) such that a recently proposed paradigm shift calls for biomechanical

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FIGURE 12.8 Peak in-shoe plantar pressure measurements during walking for an at-risk individual wearing (A) a standard insole and (B) an insole that has been optimized based on plantar pressure measurements. Note the reduction in peak pressure around the heel and forefoot regions.

evaluation as one of the important pillars for risk assessment and individual footwear optimization [28]. Further approaches, such as including finite element modeling techniques in the design process, are also being explored [29].

12.3.2 Children’s flatfoot For the first decade or so of life, most children’s feet can be characterized as flexible flatfoot, with a lowered medial arch [30]. This is thought to be due to a combination of osseus and ligamentous laxity, increased adipose tissue, and developing neuromuscular control [31]. Childhood obesity has been shown to be associated with the occurrence of flatfoot in this population [32]. It has been argued that pediatric flatfoot or developmental flatfoot are poorly defined terms that are based on unreliable static measures and in most cases do not require clinical intervention [33]. Identifying individuals who do require treatment can be challenging. Plantar pressure measurements can provide some insight into typically developing foot mechanics during childhood [34]. As children develop, their plantar pressure patterns change quickly with the skeletal development of the foot and functional developments resulting in changes in gait mechanics. As the child grows and gains more experience of walking, pessures decrease at the midfoot and shift from the medial to lateral side of the foot, and initial heel contact usually develops within the first year of independent walking [35].

12.3.3 Sports In sports, ground reaction force and plantar pressure measurements have been used to investigate, better understand, and potentially avoid the development of running-related injuries through the development of athletic footwear that is designed to reduce the impact on the human body. The topic is still of relevance since—despite the evidence from

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FIGURE 12.9 Changes in ground reaction forces between shod hindfoot strike and instructed barefoot forefoot strike running. Vertical: A; mediolateral: B. Samaan CD, Rainbow MJ, Davis IS. Reduction in ground reaction force variables with instructed barefoot running. J Sport Health Sci. 2014;3:143 151.

running biomechanics research—all efforts to provide better, that is, more protective or “healthier” running shoes has not lead to a discernable decrease of the percentage of runners suffering from such problems over the last decades [36,37]. With respect to force measurements, the parameters of interest have focused mostly on the impact characteristcs, that is, the initial phase of the ground contact. Here, the amplitude of the force peak as well as the loading rate have been used to determine the degree of loading that is imposed on the human body. Running shoes are usually expected to reduce the force amplitude and/or loading rate to provide benefit to the runner and attenuate the shock traveling through the body. This feature of a running shoe is known as cushioning and has been a design focus for athletic footwear. It has led to the development of various technologies introducing new materials and/or mechanical elements in the midsole construction. Several shoe manufacturer companies have sought their own solution by developing running shoes with their signature cushioning systems (e.g., Adidas “Boost,” Asics “Gel,” Mizuno “Wave,” Nike “Air,” On “Cloud”). However, the effectiveness largely depends on the foot strike pattern as runners can make the initial contact in the hindfoot region (so-called heel strikers) or in the mid- or forefoot (midfoot or forefoot strikers) [38]. This information can be gained from the center of pressure data of the force plates, which can be computed into the strike index (the initial contact location expressed as a percentage of foot length). Smaller values (up to 33%) indicate heel strikers, while larger values between 33% and 67% indicate midfoot strikers, and over 67% forefoot strikers [39,40]. Research has revealed that forefoot strikers tend to attenuate the initial impact and smooth the vertical ground reaction force curve thus usually reducing the loading rate as compared to heel strikers [41]. A similar or even more pronounced effect can be found in barefoot running (Fig. 12.9) which has recently been proposed as the “more natural” and therefore less harmful running modality [41]. While this might hold true for some athletes with well-developed muscular control over their lower extremity, it might be too demanding for runners with less conditioned leg muscles. For the previous application in running research, the most interesting data are contained in the vertical and the anterior-posterior components of the ground reaction force. For other sports such as soccer, tennis, basketball, and many more court sports, the mediolateral component is more important for cutting maneuvers allowing to execute rapid directional changes. This information is helpful to determine the frictional demands of such maneuvers and to optimize the outsole design to provide adequate friction [42].

12.4

Areas of future research

The use of plantar pressure and ground reaction force measurements is likely to remain a key component of many biomechanical analyses as the field continues to move forward. From a technical perspective, reducing costs without

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affecting performance is an important goal to increase the acceptance of these systems for clinical applications. Beyond this, several systems to measure regional plantar shear forces have been developed [43] but none are available commercially. The ability to reliably and accurately collect plantar shear data at the same resolution as plantar pressure data would have several potential applications; for example, it is a factor in the development in ulcers in people with diabetes [44]. In-shoe systems for long-term monitoring, particularly to alert those with an at-risk diabetic foot to overloading that may lead to damage is an active area of research and could help to reduce the incidence of plantar ulceration [45]. In the area of sports, determining links to injury risk that can be detected by patterns in ground reaction force or plantar pressure measurements is an important area of continued research.

References [1] Wolff J. The law of bone remodelling. Berlin Heidelberg: Springer-Verlag; 1986. Available from: https://doi.org/10.1007/978-3-642-71031-5. [2] Brand P. The diabetic foot. In: Davidson J, editor. Clinical diabetes mellitus: a problem orientated approach. New York: Thieme Medical Publishers; 1986. p. 376 82. [3] Beely F. Zur Mechanik des Stehens. Longenbeck’s Archiv Fu¨r Klinische Chirurgie. 1882;27:457 71. [4] Abramson E. Zur Kenntnis der Mechanik des Mittelfusses. Scand Arch Physiol 1927;51:175 234. [5] Morton DJ. Structural factors in static disorders of the foot. Am J Surg 1930;9:315 28. [6] Elftman H. A cinematic study of the distribution of pressure in the human foot. Anat Rec 1934;59:481 91. [7] Nicol K, Hennig E. Time-dependent method for measuring force distribution using a flexible mat as a capacitor. In: Komi P, editor. Biomechanics V-B. Baltimore: University Park Press; 1976. p. 433 40. [8] Telfer S, Bigham JJ. The influence of population characteristics and measurement system on barefoot plantar pressures: a systematic review and meta-regression analysis. Gait Posture 2019;67:269 76. [9] Jenkins SPR. Sports science handbook: the essential guide to kinesiology, sport and exercise science. Multi-Science Publishing; 2009. [10] Baker R. The history of gait analysis before the advent of modern computers. Gait Posture 2007;26:331 42. [11] Liedtke C, Fokkenrood SAW, Menger JT, van der Kooij H, Veltink PH. Evaluation of instrumented shoes for ambulatory assessment of ground reaction forces. Gait Posture 2007;26:39 47. [12] Hennig E, Lafortune MA. Technology and application of force, acceleration and pressure distribution measurements in biomechanics. In: Allard P, Cappozzo A, Lundberg A, Vaughan C, editors. Three-dimensional analysis of human locomotion. New York: J. Wiley & Sons; 1998. p. 109 27. [13] Hennig E. Foot pressure measurements. In: Goonetilleke E, editor. The science of footwear. New York: CRC Press, Taylor & Francis; 2012. p. 359 76. [14] Hennig E, Nicol K. Druckverteilungsmessungen. In: Banzer W, Pfeifer K, Vogt L, editors. Funktionsdiagnostik des Bewegungssystems in der Sportmedizin. Berlin, Heidelberg, New York: Springer-Verlag; 2002. p. 150 63. [15] Davis BL, Cothren RM, Quesada P, Hanson SB, Perry JE. Frequency content of normal and diabetic plantar pressure profiles: implications for the selection of transducer sizes. J Biomech 1996;29:979 83. [16] Pataky TC. Spatial resolution in plantar pressure measurement revisited. J Biomech 2012;45:2116 24. [17] Melai T, IJzerman TH, Schaper NC, de Lange TLH, Willems PJB, Meijer K, et al. Calculation of plantar pressure time integral, an alternative approach. Gait Posture 2011;34:379 83. [18] Rosenbaum D, Hautmann S, Gold M, Claes L. Effects of walking speed on plantar pressure patterns and hindfoot angular motion. Gait Posture 1994;2:191 7. [19] Rosenbaum D, Westhues M, Bosch K. Effect of gait speed changes on foot loading characteristics in children. Gait Posture 2013;38:1058 60. [20] Pataky TC, Caravaggi P, Savage R, Crompton RH. Regional peak plantar pressures are highly sensitive to region boundary definitions. J Biomech 2008;41:2772 5. [21] Pataky TC, Goulermas JY. Pedobarographic statistical parametric mapping (pSPM): a pixel-level approach to foot pressure image analysis. J Biomech 2008;41:2136 43. [22] Castro MP, Pataky TC, Sole G, Vilas-Boas JP. Pooling sexes when assessing ground reaction forces during walking: statistical parametric mapping vs traditional approach. J Biomech 2015;48:2162 5. [23] Soares DP, de Castro MP, Mendes EA, Machado L. Principal component analysis in ground reaction forces and center of pressure gait waveforms of people with transfemoral amputation. Prosthet Orthot Int 2016;40:729 38. [24] Ruhe A, Fejer R, Walker B. Center of pressure excursion as a measure of balance performance in patients with non-specific low back pain compared to healthy controls: a systematic review of the literature. Eur Spine J 2011;20:358 68. [25] Pai S, Ledoux WR. The compressive mechanical properties of diabetic and non-diabetic plantar soft tissue. J Biomech 2010;43:1754 60. [26] Lee P-Y, Kong P-W, Pua Y-H. Reliability of peak foot pressure in patients with previous diabetic foot ulceration. Gait Posture 2019;70:6 11. [27] Tsung BYS, Zhang M, Mak AFT, Wong MWN. Effectiveness of insoles on plantar pressure redistribution. J Rehabil Res Dev 2004;41:767 74.

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[28] van Netten JJ, Sacco ICN, Lavery LA, Monteiro-Soares M, Rasmussen A, Raspovic A, et al. Treatment of modifiable risk factors for foot ulceration in persons with diabetes: a systematic review. Diabetes Metab Res Rev 2020;36(Suppl. 1):e3271. [29] Telfer S, Woodburn J, Collier A, Cavanagh PR. Virtually optimized insoles for offloading the diabetic foot: a randomized crossover study. J Biomech 2017;60:157 61. [30] Uden H, Scharfbillig R, Causby R. The typically developing pediatric foot: how flat should it be? A systematic review. J Foot Ankle Res 2017;10:37. [31] Banwell HA, Paris ME, Mackintosh S, Williams CM. Pediatric flexible flat foot: how are we measuring it and are we getting it right? A systematic review. J Foot Ankle Res 2018;11:21. [32] Yan S, Zhang K, Tan G, Yang J, Liu Z. Effects of obesity on dynamic plantar pressure distribution in Chinese prepubescent children during walking. Gait & Posture 2013;37:37 42. [33] Morrison SC, McClymont J, Price C, Nester C. Time to revise our dialogue: how flat is the pediatric flatfoot? J Foot Ankle Res 2017;10:50. [34] Hennig EM, Staats A, Rosenbaum D. Plantar pressure distribution patterns of young school children in comparison to adults. Foot Ankle Int 1994;15:35 40. [35] Montagnani E, Price C, Nester C, Morrison SC. Dynamic characteristics of foot development: a narrative synthesis of plantar pressure data during infancy and childhood. Pediatr Phys Ther 2021. [36] Nigg BM, Baltich J, Hoerzer S, Enders H. Running shoes and running injuries: mythbusting and a proposal for two new paradigms: ‘preferred movement path’ and ‘comfort filter. Br J Sports Med 2015;49:1290 4. [37] Mann R, Malisoux L, Urhausen A, Meijer K, Theisen D. Plantar pressure measurements and running-related injury: a systematic review of methods and possible associations. Gait Posture 2016;47:1 9. [38] Sun X, Yang Y, Wang L, Zhang X, Fu W. Do strike patterns or shoe conditions have a predominant influence on foot loading? J Hum Kinet 2018;64:13 23. [39] Cavanagh PR, Lafortune MA. Ground reaction forces in distance running. J Biomech 1980;13:397 406. [40] Squadrone R, Rodano R, Hamill J, Preatoni E. Acute effect of different minimalist shoes on foot strike pattern and kinematics in rearfoot strikers during running. J Sports Sci 2015;33:1196 204. [41] Lieberman DE, Venkadesan M, Werbel WA, Daoud AI, D’Andrea S, Davis IS, et al. Foot strike patterns and collision forces in habitually barefoot vs shod runners. Nature. 2010;463:531 5. [42] Wannop JW, Worobets JT, Stefanyshyn DJ. Footwear traction and lower extremity joint loading. Am J Sports Med 2010;38:1221 8. [43] Stucke S, McFarland D, Goss L, Fonov S, McMillan GR, Tucker A, et al. Spatial relationships between shearing stresses and pressure on the plantar skin surface during gait. J Biomech 2012;45:619 22. [44] Yavuz M, Ersen A, Hartos J, Schwarz B, Garrett AG, Lavery LA, et al. Plantar shear stress in individuals with a history of diabetic foot ulcer: an emerging predictive marker for foot ulceration. Diabetes Care 2017;40:e14 15. [45] Najafi B, Reeves ND, Armstrong DG. Leveraging smart technologies to improve the management of diabetic foot ulcers and extend ulcer-free days in remission. Diabetes Metab Res Rev 2020;36(Suppl. 1):e3239.

Chapter 13

Electromyography and Dynamometry for Investigating the Neuromuscular Control of the Foot and Ankle Brian H. Dalton1 and Geoffrey A. Power2 1

School of Health and Exercise Sciences, University of British Columbia Okanagan, Kelowna, BC, Canada, 2Department of Human Health and

Nutritional Sciences, College of Biological Sciences, University of Guelph, Guelph, ON, Canada

Abstract The foot and ankle comprise a unique and valuable component of the body that allows humans to navigate through and interact with the surrounding environment. The foot and ankle complex is a rigid, yet active, structure for ground reaction forces to propel the body through space, overcome obstacles, and provide a base of support for stabilizing the body and maintaining upright posture. The feet and ankles also supply a rich source of sensory information for the central nervous system to interact and react to the external environment while performing activities of daily living. To extract important information regarding the neuromechanical coupling of the muscles controlling the foot and ankle, a combination of electromyography (EMG) and mechanical recordings are required. In this chapter, we discuss EMG and dynamometer experimental designs that can be used to investigate neuromuscular control with specific considerations for the intrinsic foot and ankle joint musculature. We also highlight some of the current knowledge gleaned regarding neuromuscular physiological insights using experimental designs involving EMG and dynamometry.

13.1

Introduction

The neuromuscular system is a remarkable, highly sophisticated organization of neural pathways that interpret and transform sensorimotor signals into coordinated muscle contractions for fine motor control, locomotion, and standing balance. For example, standing balance—often perceived as virtually effortless—involves complex integration and processing of sensory cues arising from the visual, vestibular, and somatosensory systems to control muscles involved in the postural task [1]. All motor signals from the central and peripheral nervous systems are transmitted through a final common pathway, the motor unit (MU), which is composed of an alpha-motor neuron and all the muscle fibers it innervates. The MU is indeed the fundamental functional unit of the neuromuscular system; whereby neural signals converge onto motor neuron pools and this neural drive is translated into muscle contractions for the production of appropriate movements. Motor output is not generated in isolation of a single MU, but via activation of populations of MUs through a combination of neural strategies known as recruitment and rate coding to produce a requisite force [2]. This force gradation depends on the quantity of active MUs, timing of the MU activity, whole muscle contractile properties, and mechanical features of the connective tissue [3]. As such, practical and sophisticated tools have been developed to assess neuromuscular activity and mechanical output, representing activation of corresponding MU pools, via EMG and dynamometry, respectively. The next sections of this chapter will briefly introduce these techniques and then provide specific technical and physiological considerations for experimental investigations related to the ankle and foot musculature.

Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00042-1 © 2023 Elsevier Inc. All rights reserved.

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Electromyography

13.2.1 Surface electromyography The investigation of human motor behavior often requires insight into the neural control strategies or amplitude of neuromuscular activation corresponding to the task of interest. Surface EMG can provide an indirect, but noninvasive, approach to assess activity of the MU pools generating the required forces to accomplish a specific motor task. The interference pattern sampled using surface EMG represents the summation of the active MUs’ muscle fiber action potentials located beneath the recording electrodes and is an indication of neuromuscular activity. As such, EMG is a relevant tool for many neuromuscular physiology, motor control/learning, and biomechanical laboratories. However, the neuromuscular activity sampled with surface EMG must be used and interpreted with caution, owing to a limited relationship with neural drive, inherent limitations, and its context-dependent nature [3,4]. Although surface electrode arrays and grids for sampling high-density surface EMG [5,6] are becoming more prominent with advances in technology and cost of specialized equipment, the basic experimental set-up involves singlechannel recordings with a pair of surface EMG electrodes positioned on the skin over the muscle belly; or an active electrode placed over the muscle belly and the other situated over the distal tendon of the agonist muscle. When using surface EMG, electrode placement (e.g., inter-electrode distance, orientation with muscle fibers) and other important guidelines should be considered to ensure adequate data acquisition. For example, surface electrodes should be positioned in line with the muscle fibers [4] with an inter-electrode distance of 20 mm center-to-center [7]. For further indepth details, excellent consensus reports regarding general guidelines for the application of surface EMG electrodes are available elsewhere [6,8]. An important consideration when interpreting surface EMG is the quantification of muscle activity for the indirect assessment of neural drive to the muscle, which is sometimes reported in absolute values of µV or mV [4]. A common procedure to estimate the magnitude of neuromuscular activation or amount of surface EMG activity is to transform the signal into a positive waveform (e.g., full wave rectification) and evaluating the mean value from the signal (Fig. 13.1) [4] or using the processed data to identify neural modulation of MU activity via correlational analyses [9,10]. Some of the most common parameters used to analyze surface EMG activity include: the average value of a full-wave rectified signal, which parallels the mean value of the absolute amplitude over a designated epoch; and root mean square (RMS) amplitude. The RMS amplitude is likely a more appropriate measure of neuromuscular activity [4], but for most common uses of surface EMG—especially those involving isometric contractions—both parameters seem adequate [4]. Other amplitude parameters are also useful and further details on these can be found elsewhere [5]. Multiple non-neurophysiological factors (e.g., EMG electrode size, inter-electrode distance, thickness of noncontractile tissue below the electrodes, movement of muscle fibers in relation to skin) can influence the surface EMG FIGURE 13.1 Example of unprocessed (top), rectified (middle), and integrated surface (bottom) EMG signals [11].

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signal and potentially complicate interpretations of unprocessed EMG signals. As such, absolute values (e.g., mV) taken from the unprocessed surface EMG are insufficient to evaluate neuromuscular activation and support neurophysiological conclusions. It is important to employ normalization procedures that can accurately represent alterations in surface EMG amplitude over time and compare differences between experimental conditions and populations (e.g., adult aging). A common normalization procedure involves reporting the surface EMG amplitude as a relative value of that collected during a maximal voluntary isometric contraction (MVC). However, the value taken from the MVC is not consistent across laboratories. For example, some researchers will take a surface EMG amplitude value for a specific epoch around the MVC peak torque/force or others will use the mean value about the peak from multiple MVC attempts [4]. Regardless of the normalization approach, comparisons between participants performing the same motor task can be afforded (e.g., comparing males and females or older and younger adults) if the MVC procedures are well controlled (see below for further details regarding MVC methodological considerations). One important consideration is whether participants achieved a “true” maximal effort during the MVC attempts. For example, some muscle groups, such as the plantarflexors [12,13] provide a greater challenge for individuals to achieve high levels of voluntary activation compared to others (e.g., dorsiflexors) [13]; whereas specific populations (older vs younger adults) may require greater practice than others to ensure a maximal voluntary effort is achieved [14]. To account for volitional effort in the normalization procedure, another means of normalizing the unprocessed surface EMG signal is to report values as a percentage of the maximal amplitude of the electrically evoked compound muscle action potential (also termed the M-wave). The M-wave represents the summation of all muscle fiber action potentials sampled by the surface EMG electrodes. Although the muscle fibers are recruited dissimilarly from a voluntary contraction, the M-wave is indicative of the maximal EMG amplitude of the muscle being recorded and provides a maximal value to normalize EMG activity. This method indeed is affected by several biological factors [15] and inherent limitations, but it does control for alterations in recording conditions [16]. Because the M-wave involves synchronous activation and amplitude cancellation will differ from voluntary actions, M-wave normalization was recently recommended to be primarily used for the normalization of electrically evoked potentials [8]. When possible, the recommendation for normalization procedures of a voluntary submaximal contraction is to use the EMG amplitude of a context-specific MVC [8].

13.2.2 Indwelling electromyography and motor unit recordings With the advances in technology, noninvasive methods to evaluate and sample discrete MU behavior during various contractions is becoming more popular and methodologies involving high-density surface EMG are quite good at sampling discrete MU action potential trains during low to high intensity isometric contractions [17,18] with a capability to track populations of specific MUs across multiple testing sessions up to days and even weeks [18]. However, some limitations still remain—especially when attempting to sample individual MU behavior during dynamic contractions and from smaller or deeper muscle groups [8]. In addition, progress is still required to improve the technology as well as the decomposition of surface EMG signals [19], but high-density surface EMG decomposition techniques are becoming a widely accepted source for MU sampling. Historically, most techniques used to sample single MU action potential trains involve placing the electrodes within the muscle of interest. One of the most commonly used intramuscular EMG techniques involves inserting fine-wire electrodes into the muscle of choice with a disposable hypodermic needle (Fig. 13.2) [20]. Once inserted within the desired muscle, the hypodermic needle is removed and the wires remain embedded in the muscle owing to the bent or “hooked” electrode tips. The electrodes are coated in an insulation and, depending on the purpose of the needle recording, a select portion of the insulation is removed to expose a small recording area at or close to the tip of each fine-wire electrode. Although this technique has been used for 60 1 years and includes some inherent limitations (e.g., signal attenuation over time [6,21], it is still a powerful tool to sample individual MU discharge properties as well as global multi-unit activity, especially in smaller muscle groups such as the intrinsic foot muscles [22,23]. A key benefit of the fine-wire technique is the ability to track discrete MU action potential trains and quantify MU discharge properties in relation to the sampled MU’s recruitment threshold. Owing to a greater EMG interference pattern with increasing contraction intensity, one fine-wire EMG electrode pair is limited to sampling only a few discrete MUs simultaneously. Hence, innovative methodologies exploiting decomposition procedures from high-density surface EMG are of great interest for many laboratories [19]. Fine-wire recordings have been employed for low intensities during controlled ramp isometric contractions [24], but have also been used successfully to sample MU properties during ramp efforts up to MVC for small [23] and large [25] muscle groups as well as during fast ballistic isometric [26] and dynamic [24] tasks. Unfortunately, the fine-wire indwelling EMG technique is not always feasible when researchers are interested in high intensity muscle actions such as an MVC.

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FIGURE 13.2 Representation of an early version of a fine-wire bi-polar indwelling electrode. These electrodes involve the insertion of the hypodermic needle into the muscle of choice. When the hypodermic needle is removed, the hooked wires, with insulation removed for recording the electrical activity, remain within the muscle [20].

The tungsten microelectrode technique affords sampling of discrete MU action potential trains during brief steady-state isometric contractions ranging from low-to-high intensities, including maximal effort. The use of tungsten microelectrodes for intramuscular MU recordings was popularized by Brenda Bigland-Ritchie and colleagues [27,28] and has been adapted by several research labs investigating MU discharge rates. When performing brief isometric contractions with the microelectrode inserted, the experimenter slowly maneuvers the microelectrode within the active muscle and once the contraction is halted, the needle is repositioned intramuscularly to sample from as many MUs within a given muscle as possible. Owing to the ease of microelectrode movement, a considerable limitation of the technique is the inability to reliably record and evaluate MU recruitment thresholds and the corresponding discharge rates at recruitment. Nevertheless, a highlight of the microelectrode technique is the sampling yield, such that, during brief isometric contractions, multiple discrete MU action potential trains can be sampled and the behavior of a population of MUs can be evaluated from one individual. The tungsten microelectrode technique indeed affords unparalleled insight into MU discharge rate behavior during highintensity contractions, which is often not achieved with fine-wire indwelling experimental procedures or other techniques. Thus, there is a trade off when selecting specific techniques. If stable MU recruitment thresholds are important at lower to moderate contraction intensities, fine-wire EMG electrodes are a suitable choice. If the research question involves evaluating MU discharge rates at high contraction intensities near maximal effort, microelectrode sampling could be the better option. Yet, it is not unwarranted to combine multiple intramuscular techniques to provide a clearer representation of MU behavior during motor control tasks [29], including high-density surface EMG.

13.3

Dynamometry

When evaluating neural control of human movement, it is important to assess not only neuromuscular activation (e.g., EMG) but also the subsequent muscle force or power generated by the MUs. In a laboratory setting, the evaluation of neuromuscular function—including force or power generating capacity—typically involves custom-made isometric myographs or commercially available multi-joint dynamometers (e.g., Biodex or HUMAC NORM Systems) that isolate muscle groups acting about a single joint. The maximal force or power produced can then be combined with other techniques used to assess neuromuscular factors underlying motor output or normalize submaximal efforts across populations (e.g., sex, age, disease-state) or conditions (e.g., fatigue, exercise-induced muscle weakness, exercise/training intervention, environmental stressor such as hypoxia, etc.). This next section will discuss the general experimental options often available when using commercially available or custom-made dynamometers: isometric, isokinetic, and isotonic.

13.3.1 Isometric A common experimental set-up for the evaluation of the mechanical output of an isolated muscle group is the use of an isometric myograph. To ensure accuracy and rigor, this testing is often confined to a single muscle group acting about a single joint while the limb and body are strapped to prevent extraneous movements and unwanted contributions of other

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FIGURE 13.3 Dynamometers for the ankle and feet. (A) Example set-up of an isometric myograph to evaluate isometric torque of the plantarflexors and dorsiflexors. The current image depicts assessment of dorsiflexion with an ankle angle of 30 plantarflexion [30]. (B) Experimental set-up to assess dynamic ankle plantarflexion and dorsiflexion using a commercially available dynamometer [31]. (C) Representation of an isometric set-up to assess force of an intrinsic foot muscle, the abductor hallucis [32].

muscles to the motor output being evaluated (Fig. 13.3). The isometric myograph incorporates a linear sensor to evaluate force of a given muscle group or a strain gage positioned at the axis of rotation to assess torque about the joint. Depending on the research question, a triaxial force sensor can also be used to evaluate muscle groups that produce force in multiple directions. When assessing neuromuscular function isometrically, investigators incorporate both involuntary (i.e., electrically evoked) and voluntary contractions or a combination of both. Involuntary isometric contractions involve electrically evoking a muscle action with a constant current stimulator and stimulating electrodes positioned over the peripheral nerve trunk or the muscle belly. Thus, the force or torque generated in response to electrical stimulation affords conclusions related to factors at or distal to the neuromuscular junction [33]. Further, some muscles provide better models of study than others, owing to the accessibility of the peripheral nerve branch, number of agonist and antagonist muscles innervated by the peripheral nerve, and quantity of muscle groups contributing to the motor task of interest. For example, the plantarflexors (tibial nerve stimulation at the popliteal fossa) and the dorsiflexors (stimulation of the deep branch of the peroneal nerve posterior to the fibular head) have accessible motor nerve branches for stimulation protocols. However, not all muscle groups with accessible nerve branches provide adequate models for electrical stimulation. The posterior tibial nerve is accessible for stimulation of the abductor hallucis, but peak torque is confounded by the activation of other intrinsic foot muscles. Therefore, superficial muscles that have inaccessible motor nerve branches can be stimulated with electrode pads placed directly over the muscle to activate their intramuscular nerve branches [34]. Caution must be taken when using electrical stimulation as the current may spread to nearby muscles when using both nerve trunk and pad stimulation.

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(B)

(A)

Single pulse

Stimulation sequence

100 Hz

10 Hz

5 Nm 500 ms

FIGURE 13.4 (A) Schematic of a sequence of electrically evoked contractions of the dorsiflexors that included a single pulse (twitch) followed by a doublet at 10 and 100 Hz and finally, a 1-second train at 10 and 100 Hz. (B) Representation of a 10 and 100 Hz doublet and 1-s train pre and post a series of maximal eccentric dorsiflexion contractions to assess exercise-induced muscle weakness [30].

One frequently used method to assess whole muscle contractile properties is the use of a single pulse—known as a twitch—or a high-frequency double pulse (typically 10 ms inter-pulse interval)—known as a doublet (Fig. 13.4). A doublet is often preferred when the signal-to-noise ratio may be an issue (e.g., fatigue, acute or chronic muscle weakness, neuromuscular disorders, etc.). The protocol involves increasing the stimulator current until a plateau in the amplitude of the evoked motor response is achieved. Then, to ensure all motor axons are contributing to the electrically evoked twitch/doublet amplitude, the intensity is increased further to B115% 150% of that needed to produce a maximal response. Thus, a supramaximal stimulation current is used to evoke maximal twitch and doublet forces during experimental paradigms [34]. These peripheral responses help investigators characterize the speed and amplitude of the whole muscle and include: peak twitch amplitude, time to peak twitch amplitude, half relaxation time as well as rate of force development and rate of relaxation. The electrically evoked parameters can provide insight into the intrinsic properties of whole muscle. For example, the half relaxation time is linked to calcium re-uptake or crossbridge detachment as a high-intensity fatigue-related increase in muscle relaxation time can be reflective of both a slowing in cross-bridge kinetics or calcium uptake [35]; whereas isometric peak torque generation is influenced by sarcoplasmic reticulum calcium handling [36]. Twitch responses are also dependent upon the activation history of the muscle and as such, are highly sensitive to potentiation [37]—a state of the muscle that is primarily characterized by an increased sensitivity to calcium via phosphorylation of myosin light chains following a high-intensity voluntary or electrically evoked contraction [37]. Thus, the interplay between fatigue and potentiation [37] must be taken into consideration when evaluating electrically evoked contractions. It is important to assess twitch and doublet parameters in a potentiated state following an MVC effort, especially when characterizing fatigue or contractile impairment of the muscle [38]. Further, peak twitch/doublet amplitude is influenced by musculo-tendinous stiffness or series compliance [39,40]. Another assessment of peripheral responses is the use of a supramaximal tetanic contraction consisting of either low or high frequencies (Fig. 13.4). The ideal presentation is to apply the electrical stimuli to the peripheral nerve for a duration (B1 second) that ensures a plateau in amplitude of the high-frequency tetanic contractions [30]. Because this technique can be prohibitively uncomfortable, induce muscle cramping, or confounded by displacement of the nerve underneath the stimulating electrode, some researchers employ high-frequencies consisting of paired pulses, despite additional limitations [30,34] such as the influence of potentiation. Often, high-frequency tetanic contractions are combined in the same protocol and compared to low-frequency tetanic contractions (e.g., 10 50 Hz). Following exercise or activity of a muscle group, especially following eccentric exercise, impairments within the muscle are dependent upon the frequency or long lasting (several days post activity) [41]. When a low-frequency is compared to a high-frequency tetanic contraction, a reduction in the ratio, known as prolonged low-frequency force depression, can be attributed to excitation contraction coupling impairments; more specifically, decreased myofibrillar calcium sensitivity and/or sarcoplasmic calcium release [42]. Alternatively, a reduction in electrically evoked force at high-frequencies is characterized by a speedy recovery, which is likely reflective of an accumulation of extra-cellular potassium [43]. The evaluation of neuromuscular function often involves the assessment of an isometric MVC. Because these maximal efforts are volitional in nature, performance is limited by supraspinal and spinal factors [44] as well as those at or below the neuromuscular junction and within the muscle [45]. Although MVCs are frequently used in experimental paradigms (e.g., training, fatigue, adult aging, etc.) and the task seems quite straightforward, considerable steps must be taken to ensure participants are highly motivated and performing maximal exertion or the influence of supraspinal factors will be magnified [44]. In an excellent and comprehensive review, Gandevia [44] outlined key experimental

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procedures that should be followed to ensure valid MVCs are performed. Briefly, the following steps should be included: (1) clear instructions and familiarization; (2) visual feedback during all maximal efforts; (3) consistent and strong verbal encouragement from the investigators; (4) participants can reject attempts that they deem submaximal; (5) when performing repeated contractions over time, the gain of the feedback should be modified so the participant is not aware of any changes in MVC amplitude; and (6) with testing over multiple sessions, provision of rewards may be considered [44]. Without consideration of the aforementioned details, voluntary activation and subsequent force output will likely be submaximal as well as variable. One crude procedure to assess whether participants are capable of achieving near maximal activation during an isometric MVC is comparing the force generated during a voluntary maximal effort with that of a supramaximal highfrequency tetanus [44,45]. A seminal study by Merton [45] using the adductor pollicis—an intrinsic hand muscle that adducts the thumb—found that torque generated by a volitional MVC was equivalent to or less than that of a 50 Hz tetanic contraction; thus, indicating that any limitations in force generating capacity is within the muscle [45]. However, comparing a volitional effort to an electrically evoked tetanus has limitations. A voluntary effort may include synergistic contributions from other muscles that may not be innervated by the motor nerve branch under investigation. For example, the peroneal muscles will contribute to voluntary plantarflexor force generation, but these muscles are not innervated by the tibial nerve, which is typically the nerve branch stimulated when evoking a plantarflexion contraction. Another technique, first introduced by Merton [45] as twitch occlusion, uses a supramaximal pulse to the corresponding peripheral motor nerve of an agonist muscle during the plateau of a maximal isometric effort. The underlying premise involves the assumption that the electrically evoked stimulus will not evoke any further increases in force if the participant is contracting the muscle with a maximal voluntary effort. Merton [45] concluded that maximal volitional output was achievable for the adductor pollicis as no additional force was evoked with the presentation of the twitch stimulus, a finding that was corroborated using other muscle models [13,14]. Thus, during an MVC, it was presumed that all MUs can be activated and driven at optimal frequencies to produce maximal force generating capacity [45]. However, low resolution of the force signal used in early studies, may have led to erroneous conclusions that central factors are not limiting the achievement of maximal force generating capacity [44]. For example, several years following Merton’s seminal study [45], Gandevia’s group [46] concluded that participants are typically unable to fully activate their thumb adductors (same muscle model used by Merton [45]), but near maximal voluntary activation is achievable in healthy individuals for many limb muscles [13,14], including those of older adults [14]. Further, the assumption of high voluntary activation attainment is based on the premise that researchers provide adequate familiarization and practice, although the values may fluctuate depending on the muscle group tested [13]. The twitch occlusion methodology has evolved since its introduction by Merton [45] with two experimental paradigms emerging: the interpolated twitch technique (ITT) and central activation ratio (CAR). Both the ITT and CAR use a supramaximal stimulation that is presented as a single, double or train of pulse(s). The CAR often involves using a train of stimuli to elicit an increase in force during the plateau of an MVC and is quantified with the following equation CAR 5 MVC/Total Force. The MVC value represents the amount of force generated from a voluntary effort whereas the Total Force value is the summation of force generated from the voluntary effort and the additional force produced by the tetanic stimulation [47]. Thus, complete voluntary activation is represented by 1.0. However, a limitation of CAR is that it compares an involuntary means of activating a muscle group to a volitional effort whereby the force generated from each stimulus is representative of different neural activation patterns. The ITT attempts to account for these dissimilarities whereby the force generated by an electrically evoked interpolated twitch (superimposed twitch) during the plateau phase of the MVC is compared to a resting, potentiated twitch, when the muscle group is relaxed fully with the equation, Voluntary Activation (%) 5 (1 2 superimposed twitch/resting twitch) 3 100 [48] (Fig. 13.5). The resting twitch is elicited following the MVC effort to ensure the muscle is in a potentiated state, which is a better representation of the active muscle fibers during the voluntary maximal effort than an unpotentiated resting twitch [48]. One caveat regarding voluntary activation is that when using electrical peripheral motor axon stimulation, additional force evoked by a superimposed twitch during an MVC can be attributed to suboptimal motor neuronal output of the muscle fibers, but the neural source of the limitation is unknown. However, transcranial magnetic stimulation can also be used to evaluate voluntary activation. If an increase in force is elicited with transcranial magnetic stimulation during an MVC attempt, the limitation in voluntary drive can not only be attributed to suboptimal motor neuron output, but also motor cortical output [48]. When assessing voluntary activation, caution should also be taken to avoid inadvertent stimulation of antagonist motor neurons and subsequent activation of the antagonist muscle group [48]. Despite the valuable insight that the assessment of the interpolated twitch can provide and its extensive use over the past several decades, there is still an ongoing debate as to its validity in providing information regarding voluntary activation of a

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Torque Resting twitch

b

a

Stim

Superimposed twitch

5 Nm

50 ms FIGURE 13.5 Visual representation of plantarflexor torque evoked from a supramaximal stimulation of the triceps surae when the muscle group is relaxed fully (b) following an MVC as well as a superimposed twitch (a) during the plateau of an MVC [48].

muscle group [49]. Regardless of the limitations, voluntary activation is a reliable tool [50] and is indeed sufficient to provide an indication of alterations in voluntary drive following specific interventions or between populations [49].

13.3.2 Dynamic In addition to isometric assessment of neuromuscular function, dynamic tasks are also commonly used owing to the quality of commercially available dynamometers (Fig. 13.3). These devices are used to isolate a single muscle group and evaluate neuromuscular parameters influencing muscle power, which is the product of the generated force or torque and the velocity of the limb through a given range of motion. Indeed, muscle power is considered an important factor contributing to overall function, especially for those with limited capacity, such as older adults [51,52]. One dynamic paradigm that has been used extensively involves isokinetic contractions, which are characterized by a constant velocity that is maintained by the dynamometer throughout a given range of motion, while the force produced is unconstrained and dependent on the effort and capacity of the participant. Because the velocity is fixed, a benefit of the isokinetic paradigm is the ability to control the duty cycle of the task. Further, for slow to moderate speed (e.g., 5 100 degrees/s) concentric and eccentric isokinetic contractions, twitch interpolation can be applied during a maximal effort to assess voluntary activation with the CAR [53]. It does seem that younger and older adults are capable of near maximal voluntary activation when performing isometric as well as concentric and eccentric isokinetic contractions [53] and thus, are not limited by voluntary drive during dynamic single limb movements, at least for those involving constant slow to moderate, fixed velocities. However, caution must be taken when characterizing isokinetic function at faster absolute velocities because as the velocity increases difficulty to “keep up” with the dynamometer increases. For example, Lanza et al. [54] reported that at speeds $ 180 and 330 degrees/s for the dorsiflexors and knee extensors, respectively, some participants, especially older adults, were unable to achieve the target velocity. Thus, the isokinetic mode may not be suitable for some testing scenarios such as evaluating rapid contractions or comparing special populations with reduced neuromuscular capacity (e.g., older adults). Another valuable paradigm for evaluating neuromuscular function during dynamic tasks is the isotonic mode of commercially available or custom-made dynamometers. Isotonic contractions represent an action whereby the velocity of the limb is unconstrained while the resistance or load is fixed. These contractions are often referred to as isotoniclike or velocity-dependent [55] as the movements are not purely isotonic owing to the braking system of the dynamometers. During isotonic-like efforts the participant is contracting against a submaximal load or resistance and as such, the evaluation of voluntary activation is currently unattainable in this paradigm. To ensure a maximal effort during the velocity-dependent concentric contractions, careful consideration must be taken and the participant’s ability to adhere to the instructions is quite important. The participant should be instructed to move the dynamometer attachment “as hard and as fast as possible throughout the entire range of motion” for each dynamic contraction [55]. Similar to an isometric MVC, adequate verbal encouragement and visual feedback of the motor output (e.g., velocity or power) should be provided [55]. Although isotonic-like contractions are relevant to real-world activities, familiarization and practice

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are necessary to ensure reproducibility [55]. Nevertheless, velocity-dependent contractions and the corresponding assessment parameters demonstrate adequate test-retest reliability [55] for various muscle groups. Both isokinetic and isotonic contractions provide valuable information regarding muscle power and the force velocity relationship [56,57], which, as mentioned earlier, is related to function [51,52]. However, other important factors that can be obtain from dynamic contractions include time-dependent variables. To produce rapid movements, ballistic rate of torque development, preceded by a rapid rise in EMG amplitude or rate of activation, is required to overcome a given resistance and initiate limb movement. Once the isometric phase is complete, rapid acceleration (e.g., rate of velocity development) is necessary to achieve peak power quickly. These time-dependent parameters may be more sensitive to acute (e.g., fatigue) and chronic (e.g., adult aging) alterations within the neuromuscular system than peak torque, velocity, or power [58,59]. For example, age-related reductions in rate of velocity development for the knee extensors are greater than peak velocity [58]. Further, Davidson et al. [60] reported that when ankle range of motion is limited, dorsiflexor rate of torque development becomes a key factor in generating peak power. Thus, neuromuscular performance may be more reliant on the time to achieve peak power rather than absolute velocity or power generation capacity per se. Nevertheless, assessment of neuromuscular function using dynamic contractions is an important evaluation paradigm.

13.4

Ankle and foot related considerations and insights

In the final section of this chapter we will provide an overview of key insights regarding dynamometry and EMG when investigating neuromuscular function of the muscles acting about the ankles and within the feet.

13.4.1 Motor unit behavior and quantity A key feature of understanding the neural control of a muscle group is evaluating the behavior of the constituent MUs. Many muscle groups of the ankle and foot are composed of predominantly type I muscle fibers [61] such as the soleus ( . 85% Type I) and tibialis anterior ( . 70% Type I). Although, the soleus and gastrocnemii comprise the triceps surae and represent the major contributors to plantarflexor torque and power [62,63], the medial and lateral gastrocnemius are made up of only B40% 50% Type I muscle fibers [61]. Similarly, the soleus exhibits slower whole muscle contractile properties than the gastrocnemii [64]. As such, the neural control strategies also vary across these muscle groups. The MU pools of muscles with slower contracting muscles achieve maximal MU discharge rates that are lower than that of the tibialis anterior. For example, the soleus exhibits mean maximal MU discharge rates of B11 17 Hz [27,65]; whereas, an intrinsic foot muscle, the flexor hallucis brevis, achieves peak MU discharge rates during a ramp isometric contraction up to maximal effort of 10 34 Hz (mean: 22 Hz) [23]. Surprisingly however, the faster contracting gastrocnemius only displays mean MU discharge rates during MVC of B21 22 Hz for both muscle components [66]. The tibialis anterior seems to express the highest MU discharge rates during a sustained maximal effort when considering the ankle and foot musculature (Table 13.1) with mean values reaching .42 Hz for younger and .30 Hz for older

TABLE 13.1 A representative overview of experimental recordings of motor unit discharge rates (MUDRs) during maximal voluntary isometric effort in humans focusing on muscles actuating the ankle or great toe, in vivo. Authors

Muscle

MUDR (Hz)

Technique

Population

Tibialis anterior

B42

Tungsten microelectrode

Young

Connelly et al. [67]

Tibialis anterior

B31

Tungsten microelectrode

Older

Bellemare et al. [27]

Soleus

11

Tungsten microelectrode

Young

Dalton et al. [65]

Soleus

17

Tungsten microelectrode

Young and older

Kirk et al. [66]

Medial Gastrocnemius

21

Tungsten microelectrode

Young and older

Kirk et al. [66]

Lateral Gastrocnemius

22

Tungsten microelectrode

Young and older

Aeles et al. [23]

Flexor Hallucis Brevis

10 34

Fine-wire

Young

Macefield et al. [28]

Extensor Hallucis Longus

17

Tungsten microelectrode

Young

Connelly et al. [67]

Values are means or a range depending on the study. a Depicts values interpreted from figure.

a a

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adults [67] adults. From an upper limb perspective, maximal MU discharge rates can reach up to 65 and 70 Hz for the biceps and triceps brachii (mean: B40 Hz) of young adults during a brief, sustained isometric effort [68]; whereas, an intrinsic hand muscle, the adductor pollicis, exhibits mean maximal MU discharge rates of 30 Hz [27]. Taken together, MU discharge rates sampled from the muscles of the ankle and foot during a brief MVC seem to be slightly lower than those of the upper limb. The neural control strategy of a given muscle is dependent upon MU rate coding and recruitment. In the upper limb, MU recruitment thresholds were observed up to 88% MVC in the biceps brachii during a ramp isometric elbow flexion. However, for the adductor pollicis, all MUs were recruited by 50% MVC with most MUs recruited within 30% MVC [69]. The narrow and low MU recruitment window within the adductor pollicis likely indicates that this intrinsic hand muscle relies preferentially on rate coding to modulate force while the biceps brachii depend on MU recruitment over a wider range of force gradation. In reference to the lower limb, MUs are recruited up to almost maximal contraction intensity for the soleus [25] and an intrinsic foot muscle that flexes the great toe [23]. Although, investigation is not exhaustive on the muscles acting about the ankle and within the foot, it seems that force production is regulated by MUs with a wide range of recruitment thresholds. These neural control strategies of recruitment and rate coding are likely owing to anatomical location and/or function of the muscle groups. The rate at which torque and velocity are produced for an isometric and dynamic task, respectively, will also influence the neural control of the muscles involved in the task, such that lower MU recruitment thresholds and faster firing rates are a function of quicker rates of torque development or velocity [17,24,70]. When compared to slow ramped increases in isometric torque that involve a progressive, concomitant increase in MU firing rates, rapid contractions involve very fast initial firing rates that subsequently decrease thereafter [26]. For example, ballistic, isometric dorsiflexion contractions are accompanied by initial MU firing rates that can exceed 200 Hz [26]. Furthermore, even minor increases in rate of torque development of the plantarflexors can shift the MU recruitment thresholds to lower values [70]. Thus, the neural control of high contractile speeds and rapid rates at which torque is generated is dependent on the ability of the central nervous system to recruit MUs quickly and achieve high initial MU firing rates [17]. In addition to MU behavior, the number of MUs within a muscle can be quantified based on indirect, but minimally invasive, assessments using electrophysiological methods, known as MU number estimations (MUNEs) [71]. MUNE techniques involve the collection of two basic size parameters of a given muscle: (1) the electrical size representing the total MU pool, which is represented by a size parameter of the M-wave and (2) the electrical size of individual MUs, which are represented by an average of the single surface MU potentials. The M-wave is collected via supramaximal peripheral nerve stimulation; whereas the surface MU potentials are determined via needle EMG with spike-triggered averaging and/or decomposition-based quantitative EMG or graded and repeated submaximal stimulation of the peripheral nerve. The MUNE is then derived by dividing the M-wave size (i.e., total size of the MU pool) by the surface MU potential size (i.e., average MU size). Thus, the MUNE is represented by the equation, MUNE 5 M-wave/surface MU potential [71]. A caution of MUNE techniques is that they likely do not represent the “true” number of MUs within a given muscle; thus, comparisons between muscles are not ideal [72]. However, MUNE assessments are reliable [73] and sensitive to age-related decrements [74,75] and neurological disruptions [76], which afford comparisons longitudinally and cross-sectionally for a given muscle group. Age-related MUNE decrements generally range from 40% 60% [72] and have been documented for muscles of the ankle and foot, which include the soleus, tibialis anterior, and extensor digitorum brevis [72]. Despite a maintenance of dorsiflexion strength, McNeil et al. [75] reported lower MUNE for older adults in their seventh decade compared to younger adults. Beyond eighty years of age, dorsiflexion strength was lower compared to younger adults, which may have been influenced by an accelerated loss of functioning MUs as MUNEs were even lower in the very old adults compared to those in their sixties [75]. However, not all muscles depict typical age-related reductions. One report found that soleus MUNEs were not significantly different for individuals in their eighth decade of life compared to a young adult cohort [12]; whereas another reported a 70% lower MUNE of the soleus in males greater than 90 years than younger adults [74]. Taken together, these soleus MUNE studies indicate that MU reductions may be delayed in a muscle that is composed of predominantly slow type muscle fibers or habitually active. In support of habitual activity, Power et al. [77] reported that tibialis anterior MUNEs were not different in lifelong competitive runners compared with younger adults. In a follow up study [78], the researchers reported that tibialis anterior MUNEs were greater in world-class master’s athletes than age-matched controls, which indicated that MU quantity was better maintained in highly active master’s athletes than recreationally active older adults. Yet, this result is not always reported for master’s athletes [71].

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13.4.2 Maximal voluntary contractions and knee angle As noted earlier, the foundation of neuromuscular function is normalizing one’s effort to the maximal capacity of the specific muscle group. Typically, isometric and isotonic-like contractions are normalized to the isometric MVC. For the ankle musculature, and more specifically the dorsiflexors, high voluntary activation levels ( . 95%) are achievable in most participants with limited practice in healthy young adults [55,67], master’s athletes [78], and recreationally active old and very old adults [67,75]. On the other hand, the plantarflexors [12,13] and abductors of the great toe [79] are more difficult to activate during a maximal effort contraction compared to the dorsiflexors. However, at least for the plantarflexors, given appropriate practice and familiarization, younger and older participants are capable of high voluntary action levels [65]. An important consideration when evaluating maximal torque output of any muscle group is controlling for the torque-length relationship. For example, dorsiflexion MVC torque increased successively over an ankle angle range from 20 degrees dorsiflexion to 10 degrees plantarflexion [80]; whereas plantarflexion MVC torque progressively increased from an ankle angle positioned at 30 degrees plantarflexion to 10 degrees dorsiflexion with optimal torque generation achieved for all participants at or greater than 15 degrees dorsiflexion [81]. Modifying the ankle angle alters the length of all muscles acting about the ankle joint. However, the gastrocnemii are major contributors to plantarflexor torque and, anatomically, their tendons cross both the knee and ankle joints. When the knee is in an extended knee position, all plantarflexors contribute to overall ankle extensor torque; whereas in a flexed knee position (B90 degrees or greater), the gastrocnemii are placed at a neuromechanical disadvantage, which reduces isometric and isokinetic plantarflexor MVC torque [56,62]. Because of the limitations imposed on the gastrocnemii via knee flexion, the soleus becomes the main contributor to plantarflexor torque generation. Thus, researchers use a flexed knee position when interested in the soleus as a model to study neuromuscular function. When the knee is flexed, velocity-dependent plantarflexor peak power is reduced over a range of relative resistances from 15 to 75% MVC compared with an extended knee, but peak velocity is only reduced at low to moderate resistances from 15 to 30% MVC [56]. The reductions in plantarflexion MVC torque and power generating capacity is likely owing to the neuromechanical disadvantage placed on the gastrocnemii, but likely not voluntary drive to the plantarflexors as individuals are capable of near maximal voluntary activation ( . 95%) with both a flexed and extended knee [56]. Flexing the knee shortens the muscle fascicle length of the gastrocnemii [82] and in turn places this muscle group at a mechanical disadvantage. Hence, the gastrocnemii provide a limited contribution to overall plantarflexor output with the knee flexed and hence, electrically-evoked peak twitch torque is reduced and whole muscle contractile speed is slowed [56] compared to extended knee angle. In addition, the neural control of the gastrocnemii is also altered. For knee flexion, Cresswell et al. [62] reported that MVC torque reductions were accompanied by lower M-wave amplitudes for the medial and lateral gastrocnemius, but M-wave amplitudes were not different for the soleus between knee joint angles. Thus, indicating that knee flexion-related decrements in plantarflexion function are likely, in part, owing to decreased sarcolemmal excitability or neuromuscular propagation [62]. Further, Hali et al. [83], evaluated MU behavior of all three components of the triceps surae during ramp isometric contractions with a flexed and extended knee. Recruitment thresholds shifted towards higher levels in the flexed than extended knee for both gastrocnemii with a slight reduction in initial MU discharge rates for the medial, but not lateral gastrocnemius (Fig. 13.6). Consistently, no differences were detected for the soleus MU recruitment thresholds nor MU discharge rates across knee angles [83]. Thus, MU recruitment of the gastrocnemii as well as medial gastrocnemius MU discharge rates are inhibited in a flexed compared with extended knee position [83]. This altered MU behavior in combination with reduced sarcolemmal excitability or neuromuscular propagation and an increased mechanical disadvantage of the gastrocnemii are likely responsible for reductions in plantarflexor torque and power generating capacity during knee flexion.

13.4.3 History-dependence of force To assess voluntary or electrically evoked torque about the ankle, isometric contractions are often performed. While the torque-angle relationship and cross-bridge theory can, with great accuracy, predict an individual’s isometric torque production for a given muscle length, the theory’s predictive power falls short when the isometric contraction is preceded by an active lengthening (i.e., eccentric contraction) or shortening (i.e., concentric contraction) action [84]. This history-dependence of force was first identified almost 70 years ago, and since then it has been observed from the single sarcomere up to electrically evoked and voluntary contraction in humans (Table 13.2).

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FIGURE 13.6 (A) Motor unit (MU) properties are shown for the medial gastrocnemius (MG), lateral gastrocnemius (LG) and soleus (SOL) in an extended and flexed knee joint position (x axis). Each black dot represents an individual MU. The y axis represents the MU recruitment threshold (RT). The horizontal lines depict median values for each group. (B) The y axis represents the MU firing rates (FRs) with the horizontal lines representing mean values for each group. An asterisk represents a significant difference compared to the extended knee joint position (P , .05). (C) Each row displays an overlay of action potentials for those MUs sampled during plantarflexion contractions in both an extended and flexed knee joint position. The MU action potential overlays include 50 500 spikes per template. Six, four and six MUs were sampled from the MG, LG and SOL, respectively. (D) Joined dots represent the same MU in the extended and flexed knee joint positions. The MU RTs are depicted as percent maximal voluntary contraction (%MVC; left column) and MU FRs are highlighted in the right column in Hz [83].

The history-dependence of force is characterized by an increase (residual torque enhancement; rTE) or decrease (residual torque depression; rTD) in steady-state isometric torque following active muscle lengthening or shortening, respectively, as compared to a purely isometric reference contraction at the same muscle length and level of activation [84]. While the underlying mechanisms of the history-dependence of force are not yet completely understood, the phenomena of rTE and rTD have been well-characterized [103,104]. The magnitude of rTE is dependent upon the amplitude of muscle stretch and independent of lengthening velocity [103,104]. Consequently, owing to the increased proportion of passive force to total force production in the rTE state, lower MU activity (as measured with EMG) is required to achieve an equivalent isometric torque output (activation reduction) [103,104]. Specifically, using fine-wire EMG electrodes inserted into the tibialis anterior, it was demonstrated that the activation reduction in the rTE state was owing to less active MUs and lower MU discharge rates than a purely isometric reference contraction [102]. Following active shortening, the magnitude of rTD appears to be strongly and positively related to the amount of work (i.e., product of force and muscle length change) performed during active shortening [105]. As a result of shorteninginduced rTD, greater MU activity is required (activation increase) to offset the deficits in torque and achieve a similar isometric torque output [106]. During human in-vivo studies, rTE values have been reported to range from 5% 25% for the dorsiflexors and plantarflexors (Table 13.2); during maximal and submaximal voluntary contractions. Studies that matched force/torque reported 3% 27% decreases in activation (i.e., AR), in the rTE state (Table 13.2). Meanwhile, rTD during voluntary contractions have been reported to be 8% 39% for the dorsiflexors and plantarflexors (Table 13.2) with 9% 48% greater activation (i.e., AI) in the rTD state during torque-matched conditions (Table 13.2). The everyday implications of the history-dependence of force are continuing to be elucidated, but one clear message emerging from the literature is that this basic intrinsic property of muscle has the potential to alter feedback from the periphery to the central nervous system, which can in turn alter an individual’s voluntary motor control [89,98]. Recent studies reported that voluntary torque production in an rTE and rTD state is likely mediated by torque-dependent afferent feedback, such as a modification of neural activity via Ib afferents [100,107]. As the history-dependence of force

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TABLE 13.2 A representative overview of experimental work in humans focusing on rFD and rFE about the ankle in humans, in vivo. Authors

Muscle action

rFD/AI (%)

Approach

Settings

Tilp et al. [85]

Dorsiflexion

9 20

Voluntary activation

100% MVC

Power et al. [86]

Dorsiflexion

16 39

Voluntary activation

100% MVC

Fukutani et al. [87]

Plantarflexion

8 13

Electrical activation

50 Hz stimulation at 25% MVC

Grant et al. [88]

Dorsiflexion

15 16

Voluntary activation

100% MVC

Sypkes et al. [89]

Dorsiflexion

13

Voluntary activation

Activation matched at 40% MVC

Paquin and Power [90]

Dorsiflexion

10 17/23 48

Voluntary activation

Torque and activation matched at 20% 80% MVC

Hahn and Riedel [91]

Plantarflexion

12

Electrical activation

20 Hz stimulation to 30% MVC

Chen and Power [31]

Dorsiflexion

16 21/13 24

Voluntary activation

Maximal and submaximal

Pinniger and Cresswell [92]

Plantarflexion

13

Electrical activation

50 Hz tetanic submaximal stimulation to 15% MVC

Pinniger and Cresswell [92]

Plantarflexion

7

Voluntary activation

Activation matched at 25% MVC

Pinniger and Cresswell [92]

Dorsiflexion

12

Voluntary activation

Activation matched at 25% MVC

Tilp et al. [85]

Dorsiflexion

8 16

Voluntary activation

100% MVC

Power et al. [93]

Dorsiflexion

10 25

Voluntary activation

100% MVC

Hahn et al. [94]

Plantarflexion

9

Voluntary activation

100% MVC

Power et al. [95]

Dorsiflexion

8; 22

Voluntary activation

100% MVC

rFE/AR (%)

Fukutani et al. [96]

Plantarflexion

5 16

Electrical activation

25% MVC

Hahn and Riedel [91]

Plantarflexion

5

Electrical activation

30% MVC

Dalton et al. [97]

Plantarflexion

15/5 7

Voluntary activation

Maximal and submaximal

Sypkes et al. [98]

Dorsiflexion

10

Voluntary activation

Activation matched at 40% MVC

Chen and Power [31]

Dorsiflexion

11 13/3 12

Voluntary activation

Maximal and submaximal

Mazara et al. [99]

Plantarflexion

5 9/9 11

Voluntary activation

Torque and activation Matched at 20% and 60% MVC

Contento et al. [100]

Dorsiflexion

15/14

Voluntary activation

Torque and activation matched at 40% MVC

Marion and Power [101]

Dorsiflexion

27

Voluntary activation

Torque matched at 60% MVC

Jakobi et al. [102]

Dorsiflexion

24 44

Voluntary activation

Torque matched at 10 and 20% MVC

Values are means. AI, activation increase; AR, activation reduction; MVC, maximum voluntary isometric contraction; rFD, residual force depression; rFE, residual force enhancement.

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can influence voluntary control of force, in studies involving tasks of daily living, such as dynamic contractions, rTE and rTD should not be ignored.

13.5

Future research

The musculature acting about the ankle and within the foot provide important and valuable roles for torque and power generation to maintain upright posture and prevent the body from toppling as well as propel the body through its environment. As such, the neuromechanical interplay between these muscle groups are critical for activities of daily living. The evaluation of the torque and power generating capacity as well as the neural control of the muscle actuators acting about the ankle and within the foot are important in furthering our understanding of the underlying factors related to human movement. Indeed, a combination of dynamometry, EMG techniques, and other electrophysiological methodologies will play a key role in future research. For example, with the emergence and increasing availability of high-density EMG and corresponding decomposition software, the ability to track and sample large populations of MUs within a given muscle or from multiple muscles using various contraction types, interventions, and time durations are becoming more practical. Historically, research has been focused on an isometric model to explore MU behavior and neuromuscular function. A concerted effort is needed to fill in the gaps and expand on what has been gleaned from studies investigating isometric tasks. Because the task and criterion measure are important, an emphasis should be placed on furthering our knowledge using dynamic contractions and more specifically, isotonic-like contractions to answer questions related to the basic neuromuscular function of the ankle and intrinsic foot musculature as well as those related to activities of daily living. Much of the data to date have been focused on male participants or healthy younger adults. Thus, an emphasis should be placed on evaluating neuromuscular function in female populations and other under-represented groups to present a more wholistic understanding of the neuromechanical control of the ankle and foot. In general, little is known about the contributions of the intrinsic foot muscles to various tasks and these muscles often go overlooked, although studies involving an intrinsic foot muscle model have provided valuable insight into MU behavior [23] and the neural control of standing [22]. Future work should continue to expand upon the limited data and help elucidate the potentially complex control of the interactions between the intrinsic and extrinsic foot muscles contributing to human movement. For example, combining EMG, dynamometry, and electrophysiological techniques such as peripheral nerve or muscle belly stimulation or sensory stimulation (e.g., electrical vestibular stimulation) could help determine the contribution of various inputs to modulate MUs involved in actuating and supporting the ankle and foot. Overall, as technology advances and experimental paradigms become more advanced, our understanding of the ankle and foot from previous experiments using EMG and dynamometry will continue to grow.

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Reciprocal activation of gastrocnemius and soleus motor units is associated with fascicle length change during knee flexion. Physiol Rep 2014;2:e12044. Available from: https://doi.org/10.14814/phy2.12044. [83] Hali K, Kirk EA, Rice CL. Effect of knee joint position on triceps surae motor unit recruitment and firing rates. Exp Brain Res 2019;237:2345 52. Available from: https://doi.org/10.1007/s00221-019-05570-7. [84] Herzog W. History dependence of skeletal muscle force production: implications for movement control. Hum Mov Sci 2004;23:591 604. Available from: https://doi.org/10.1016/j.humov.2004.10.003. [85] Tilp M, Steib S, Herzog W. Force-time history effects in voluntary contractions of human tibialis anterior. Eur J Appl Physiol 2009;106:159 66. Available from: https://doi.org/10.1007/s00421-009-1006-9. [86] Power GA, Makrakos DP, Stevens DE, Herzog W, Rice CL, Vandervoort AA. Shortening-induced torque depression in old men: implications for age-related power loss. Exp Gerontol 2014;57:75 80. Available from: https://doi.org/10.1016/j.exger.2014.05.004. [87] Fukutani A, Misaki J, Isaka T. Force depression in plantar flexors exists equally in plantar flexed and dorsiflexed regions. Front Physiol 2017;8. Available from: https://doi.org/10.3389/fphys.2017.00183. [88] Grant J, McNeil CJ, Bent LR, Power GA. Torque depression following active shortening is associated with a modulation of cortical and spinal excitation: a history-dependent study. Physiol Rep 2017;5:1 10. Available from: https://doi.org/10.14814/phy2.13367. [89] Sypkes CT, Kozlowski B, Grant J, Bent LR, McNeil CJ, Power GA. Spinal excitability is increased in the torque-depressed isometric steady state following active muscle shortening. R Soc Open Sci 2017;4. Available from: https://doi.org/10.1098/rsos.171101. [90] Paquin J, Power GA. History dependence of the EMG-torque relationship. J Electromyogr Kinesiol 2018;41:109 15. Available from: https:// doi.org/10.1016/j.jelekin.2018.05.005. [91] Hahn D, Riedel TN. Residual force enhancement contributes to increased performance during stretch-shortening cycles of human plantar flexor muscles in vivo. J Biomech 2018;77:190 3. Available from: https://doi.org/10.1016/j.jbiomech.2018.06.003. [92] Pinniger GJ, Cresswell AG. Residual force enhancement after lengthening is present during submaximal plantar flexion and dorsiflexion actions in humans. J Appl Physiol (1985) 2007;102:18 25. Available from: https://doi.org/10.1152/japplphysiol.00565.2006. [93] Power GA, Rice CL, Vandervoort AA. Residual force enhancement following eccentric induced muscle damage. J Biomech 2012;45:1835 41. Available from: https://doi.org/10.1016/j.jbiomech.2012.04.006. [94] Hahn D, Hoffman BW, Carroll TJ, Cresswell AG. Cortical and spinal excitability during and after lengthening contractions of the human plantar flexor muscles performed with maximal voluntary effort. PLoS ONE 2012;7:e49907. Available from: https://doi.org/10.1371/journal.pone.0049907. [95] Power GA, Rice CL, Vandervoort AA. Increased residual force enhancement in older adults is associated with a maintenance of eccentric strength. PLoS ONE 2012;7. Available from: https://doi.org/10.1371/journal.pone.0048044. [96] Fukutani A, Misaki J, Isaka T. Influence of joint angle on residual force enhancement in human plantar flexors. Front Physiol 2017;8. Available from: https://doi.org/10.3389/fphys.2017.00234. [97] Dalton BH, Contento VS, Power GA. Residual force enhancement during submaximal and maximal effort contractions of the plantar flexors across knee angle. J Biomech 2018;78:70 6. Available from: https://doi.org/10.1016/j.jbiomech.2018.07.019. [98] Sypkes CT, Kozlowski BJ, Grant J, Bent LR, McNeil CJ, Power GA. The influence of residual force enhancement on spinal and supraspinal excitability. PeerJ 2018;2018:1 15. Available from: https://doi.org/10.7717/peerj.5421.

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[99] Mazara N, Hess AJ, Chen J, Power GA. Activation reduction following an eccentric contraction impairs torque steadiness in the isometric steady-state. J Sport Health Sci 2018;7:310 17. Available from: https://doi.org/10.1016/j.jshs.2018.05.001. [100] Contento VS, Dalton BH, Power GA. The inhibitory tendon-evoked reflex is increased in the torque-enhanced state following active lengthening compared to a purely isometric contraction. Brain Sci 2020;10:1 11. Available from: https://doi.org/10.3390/brainsci10010013. [101] Marion R, Power GA. Residual force enhancement due to active muscle lengthening allows similar reductions in neuromuscular activation during position- and force-control tasks. J Sport Health Sci 2020;00:1 7. Available from: https://doi.org/10.1016/j.jshs.2020.07.003. [102] Jakobi JM, Kuzyk SL, McNeil CJ, Dalton BH, Power GA. Motor unit contributions to activation reduction and torque steadiness following active lengthening: a study of residual torque enhancement. J Neurophysiol 2020;123:2209 16. Available from: https://doi.org/10.1152/ jn.00394.2019. [103] Seiberl W, Power GA, Hahn D. Residual force enhancement in humans: current evidence and unresolved issues. J Electromyogr Kinesiol 2015;25:571 80. Available from: https://doi.org/10.1016/j.jelekin.2015.04.011. [104] Chapman N, Whitting J, Broadbent S, Crowley-McHattan Z, Meir R. Residual force enhancement in humans: a systematic review. J Appl Biomech 2018;34:240 8. Available from: https://doi.org/10.1123/jab.2017-0234. [105] Herzog W, Leonard TR, Wu JZ. The relationship between force depression following shortening and mechanical work in skeletal muscle. J Biomech 2000;33:659 68. Available from: https://doi.org/10.1016/S0021-9290(00)00008-7. [106] Chen J, Hahn D, Power GA. Shortening-induced residual force depression in humans. J Appl Physiol 2019;126:1066 73. Available from: https://doi.org/10.1152/japplphysiol.00931.2018. [107] Sypkes CT, Dalton BH, Stuart J, Power GA. Inhibitory tendon-evoked reflex is attenuated in the torque-depressed isometric steady-state following active shortening. Appl Physiol, Nutr, Metab 2020;45:601 5. Available from: https://doi.org/10.1139/apnm-2019-0321.

Chapter 14

From Impossible to Unnoticed: Wearable Technologies and The Miniaturization of Grand Science Eric Rombokas and David Boe Rombolabs, rombolabs.github.io

Abstract The scientific findings and the instruments used to demonstrate them follow a natural arc from exotic to pedestrian. This chapter provides an overview of historical examples in certain domains that are either currently, or soon to be, wearable on humans as a part of laboratory study or everyday life. It also provides a nonexhaustive overview of wearable products and technologies relevant to foot and ankle biomechanics.

14.1

Introduction

The present is stubbornly unremarkable. As you read this, you’re likely surrounded by technology that would bewilder and amaze your ancestors yet somehow manages to seem pedestrian, even clunky, and inadequate, to you. It has become cliche to point out the eyebrow-singeing pace of technological advancement, but in this chapter, we focus on an advance that is often overlooked: wearable technology. Improvement in machine capability has two sides. The first is the widening envelope of what is possible. Something that has previously been impossible for humans, like flying, is rendered possible through new methods or devices like the Wright flyer. The second is increasing access to that capability. Tech that begins as an exotic and elaborate “mad science” project becomes “mere science,” then a decadent optional luxury, and eventually a commonplace and necessary requirement of modern living. In the space of a single century, flying machines traversed the full range, from impossible to banal. The natural consequence of increasing access is that most technologies, however arcane their beginnings, will eventually become wearable. There exist wearable flying machines today, though they currently seem exotic and impractical. As technologies miniaturize, they waste less power, generate less heat, cost less money, and eventually become viable as worn gear, perhaps on their way to being implantable. Time will tell, though admittedly the idea of an implantable flying machine seems to strain the example. That’s kind of the point; we hope that this chapter encourages the reader to boldly consider the implications of advanced wearable technology. We will attempt to remain relevant to the foot and ankle.

14.1.1 Tech affords understanding The scholars of the world, scratching their chins and performing deductive reasoning, are ultimately limited by the availability of observations. Without measurements, examples, and data generated by the actual world, it is difficult to reach conclusions that successfully explain natural phenomena. Despite the timeless allure of theoretical rigor and logic, and the persistent idea of the triumph of human intelligence, the record is in favor of getting one’s hands dirty and Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00041-X © 2023 Elsevier Inc. All rights reserved.

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measuring real things. Most progress has been made by someone dabbling with new things without fully understanding them. With a few notable exceptions, it has been the development of technologies that enable scientific understanding, not vice versa. To restate provocatively, the history of increasing human understanding has not been driven directly by human deductive reasoning, but by new technologies that produce mysterious new capabilities that serve as a foundation for inductive reasoning. Gadgets and techniques, poorly understood at first, drive understanding, not the other way around! This is antithetical to the persistent, and pernicious, idea that smart people invent practical things. Practical people invent things and then later, smart people explain what’s happening. An example from history is microscopy. Antonie van Leeuwenhoek was a cloth merchant who became interested in lenses. He began using them to observe the threads of the cloth under magnification, and developed a means to make tiny spheres of glass for this purpose. He began looking at all sorts of things under previously impossible magnification, and eventually discovering what we would now call microbes. These activities and observations were entirely driven by practical curiosity, and the affordance of the new magnifying lenses. His findings were largely resisted and questioned by the august Royal Society scientists, theorists, and experts of human reason, until they became extensively verified and impossible to ignore. Despite millenia of conjecture and logical deductive reasoning, it was the Tech that led to understanding. This means that as technologies trend toward wearable, we will gain previously impossible insight into the secret world of the body. What knowledge might lurk in the currently immeasurable ocean of body movements, forces, sounds, strains, of every animal in the world? Might ankle sprains be foreseeable weeks in the future, in the hidden stiffnesses of ligaments or the changing tone of muscles? How cyclic are people’s movements in real life? What are the differences between how the young and the old, the swimmer and the weightlifter, negotiate the terrains of their daily lives? These questions could be answerable if everyone wore a (currently impossible) unobtrusive suite of sensors every moment of their lives.

14.1.2 Wearable domains There is limited utility for writing this book chapter as a simple catalog of current wearable devices. The pace of technology development would render it instantly obsolete. At the same time, it is important to provide an informative survey of what kinds of devices exist and an attempt to predict trends (Fig. 14.1). Therefore this chapter begins with two brief histories as a retrospective look at how technologies arise and transform. We also present an incomplete overview of the history of force and pressure, and health and activity monitoring. This will be followed by a review of the present, categorizing any wearable into one of five domains. Each domain is a class of technologies that relate to a kind of measurement or an intended use case. We have chosen Force and Pressure, Health and Activity Monitoring, Actuation and Assistance, Haptics, and Motion Capture. Arguments could be made for more or fewer than five, or other categorizations, but we feel that these align well with the way these technologies are researched, commercialized, and experienced. The reader who is only interested in becoming familiar with the current “state-of-the-art” can constrain themselves to The Present portion of the chapter. We will also briefly discuss future developments that may occur in this area. We invite you to join us for a quick tour of time first.

14.1.3 Breakout box: what makes a successful wearable device? Like other early technologies, wearables have been frequently beset with failure. Consider Google Glass: a technological marvel built by brilliant engineers, yet regrettably useless in most hands. It simply did not address a common need for enough people to be a commercially viable product. A successful wearable first and foremost solves a problem. Claude Shannon and Edward Thorp understood this when they designed and built the first wearable computer in 1961. It consisted of a pair of switches built into their shoes, a 12-transistor worn computer, and an early version of the earbud. Their grand purpose? Win big at roulette in Vegas. They succeeded a little too much, and it seems it is with pride that Thorp later wrote, “The descendants of the first wearable computer were formidable enough to be outlawed.” The simple act of donning and doffing a device on a daily basis is sufficient cause to abandon it. Human laziness cannot be underestimated when designing a wearable system. For this reason, textile-integrated electronics are promising—few people are lazy enough to forgo wearing clothes. On the other hand, if the wearable provides sufficient value, then the effort may be merited. Orthotic devices that enable someone to stand up and walk to the

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FIGURE 14.1 Author Rombokas wearing lab-based optical motion capture markers (left) and XSens motion capture units, custom force sensing boot outsole, vibrotactile sensory substitution array (right).

bathroom are well worth the trouble. Navigating this tradeoff requires careful consideration of the user and their problems.

14.2

The past

The intent of this section is not necessarily to provide a broad education in the history of science. Instead we pluck choice examples of now-familiar technology from the dusty branches of time to demonstrate how outlandish, how elaborate they would have seemed to contemporaries. Their inventors would be shocked to see you wearing them. Looking backwards, it is shocking how circuitous the path has been to the development of our modern wearables.

14.2.1 Force and pressure We humans had a slow journey figuring out how force works, which is surprising given how fundamental it is to our daily lives. Simple machines employing pulleys, levers, and gears were extensively used in antiquity despite the lack of a foundational understanding of force. One major problem was that in daily life, when one stops actively pushing something it seems to come to a halt. Motion seems to require the continuous exertion of force. How, then, does a thrown object, or an arrow, keep flying through the air despite there being nothing that continues to push it? This misunderstanding was not corrected until Renaissance figures like Galileo and Newton set things straight, accounting for friction, mass, acceleration, gravity and all of the concepts the modern reader would recognize. Nevertheless, our brains stubbornly persist with a “naive physics” that we unconsciously use to navigate daily life. This is a suite of mental simulation tools that appear early enough in life that they must have innate components. These can be tested even in preverbal infants by observing what they choose to look at—babies will gaze quizzically at physical scenarios that seem to defy physics. Heuristic physics rules, that are sometimes a little wrong [1], but useful simplifications, appear to come preloaded in human brains or to develop very early in life [2,3]. Some examples include: (1) falling objects fall straight down, and at a constant rate. (2) Objects do not pass through each other. (3) When two objects interact, one is the

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FIGURE 14.2 The 1644 experiment of Evangelista Torricelli, demonstrating the forces exerted by vacuum, air pressure, and water pressure.

“pusher” and the other one “gets pushed.” (4) The initial force imparted to an object persists as a residual force, perhaps decreasing slowly, after a projectile is flying. As a result of intuitive physics reasoning taking shortcuts like these, thought experiments are an unreliable way to reason about forces. Early instruments used to measure or reason about force and pressure include simple ramps, used by Galileo to study gravity, or the hydrostatic experiments of Blaise Pascal, inventing the hydraulic press and the syringe. These were Grand Science, occupying enormous space and involving ladders, towers, or special installations (Fig. 14.2). Arguably the first practical instrument intended for regular use was the aneroid barometer, invented in 1844 by Lucien Vidi, and distinguished by not using a liquid medium like water or mercury. It is compact and reliable, and still widely used today. It has only recently been overtaken by miniscule electronic barometers included in cellphones. A barometer is a force sensor, but not especially useful for anything but atmospheric pressure. The first broadly applicable force sensor that could be affixed to a variety of objects was perhaps the 1936 electric strain gage created by Charles M. Kearns, used to measure strain in aircraft propellers. Electric strain gages have since become highly sensitive, reliable, and miniaturized, sensing multiple axes simultaneously (e.g. Fig. 14.3). As society has industrialized, mechanical force sensing and creation have proliferated into every aspect of human life, and toward wearable applications. Mechanized industrial production has allowed for force generators—motors, speakers, hydraulics, and other actuators—to be a part of handheld tools, kitchens, cars, factories, and increasingly, our bodies. Wearable robots are beginning to transition from research prototypes into real-world applications [4,5]. These range from assistive devices for rehabilitation, for daily use to compensate for impairment, and even for augmentation of typical strength or dexterity [6,7].

14.2.2 Health and activity sensing Most people living today, even if they are not experts, have a basic knowledge of the body and how it moves. It is generally understood that muscles contract to pull bones and that nerves send signals back and forth from the extremities to

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FIGURE 14.3 Electric strain gages in a “Wheatstone bridge” circuit measuring strain at a specific mechanically flexible portion of a device.

the spine and brain. Without specific education, simply from being immersed in a modern culture, you may know details about the physiology of how that is achieved: electricity is involved, ions are exchanged, energy is released from special molecules, oxygen and carbon dioxide are created or consumed. This background knowledge makes it hard to put oneself into the mindset of the scholars of Antiquity, whose understanding of physiology was, to the modern eye, spectacularly spotty. Aristotle and compatriots raised the issue of how animals and people move and managed to barely even mention muscles at all [8,9]. Blood was theorized to be created in the heart, and used as a sort of pneumatic signaling medium, along with air (pneuma) to effect perception. A few generations later, Galen (Aelius Galenus) put forth a more modern account that set the standard for centuries. It included pulmonary, but not systemic, circulation, and posited that blood is created in the liver and then consumed by the rest of the body, rather than used as a transport medium. With some exceptions that didn’t penetrate into Western practice [10], this incomplete and incorrect model was dominant for a millennium. Contemplate a thousand years of people living, moving, being injured and recovering, butchering and eating animals, with apparently no improvement to the state of knowledge. At last, in the 16th century, the valves of the heart and veins, the two distinct loops of pulmonary and systemic circulation, and the heart as a pump instead of a generator of blood, were articulated [11]. In the 1780s, Luigi Galvani and others performed a series of experiments involving electricity and frogs and began to understand how nerves and muscles operate. Electrical stimulation of frog leg muscles causes contractions and convulsions. This and other experiments led to what we would now recognize as neuromotor physiology; nerves were not pipes for fluids, but conductors of electricity! It was experimentation and observation, not the hypothetical theorizing of the ancients, that lead to these breakthroughs. Although these early experiments were still definitely “grand science” involving elaborate room-scale apparatus, progress was steadily made throughout the years alongside increasing understanding of electricity in general. By the 20th century there had emerged practical electrophysiological uses for amplifiers, speakers, and other means of measuring and understanding the electrical biological processes of our nerves and muscles [12]. Progress in harnessing these technologies is accelerating at a dizzying pace. Affordable, reliable consumer electronics are worn by millions of people to sense the activity of their hearts, their muscles, and the movement of their bodies. Home scales use bioimpedance analysis to estimate body composition. Sophisticated sensors are used in clinical settings to detect minute variations in cardiac function and nerve conduction. Muscle activity sensors and stimulators can be placed unobtrusively onto the skin like a temporary tattoo [13,14]. Implanted pacemakers [15] and other neural stimulation systems have been in use for decades, and can serve double duty as sensors [16]. There are fabric [17] or even chronically implanted electromyographic sensors [18] suitable for control of motorized prosthetic limbs or for functional electrical stimulation [19], in which muscles are stimulated to create desirable movements of the body. These innovations are eroding the barriers between humankind and our machines. Our clothes and tools are simple ways to augment the mechanical capabilities of our bodies. These new devices will transform our bodies’ informational capabilities, and may eventually seem as prosaic, as normal and necessary, as do your shoes.

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The present

In this section, we provide an overview of the current state of the art for wearable technologies, with special attention to the foot and ankle.

14.3.1 Force and pressure sensing Although there have been recent advances in wearable devices that generate forces [7], the primary use of wearables for force remains in sensing. Forces underfoot play an outsized role in our health, and are the reason so many orthotists make custom insoles to guide the flow of forces from the ground all the way into the spine. A well-designed insole can eliminate low back pain in some cases. A poorly designed insole very well may cause it. The orthotist uses a mix of knowledge sources to accomplish this—palpation of the foot, observation of the body in static and dynamic postures, and the application of biomechanics. This is not to say that orthotists are the only experts of the foot, but with the lack of evidence to guide clinical decision-making, they substitute many years of experience and training to understand how forces act on the foot. It is complicated, and there is no guarantee of optimality. The leap from traditional, lab-based force sensing systems to wearable force sensing systems is beginning to change that. As it turns out, “force” can refer to many things, and sensing it is not as trivial as one might think. In modern times, biomechanics labs are often fitted with force plates, or panels in the floor equipped with strain gages that have been carefully calibrated. When the plate is stepped on, it reads the “ground reaction force” or the equal and opposite force generated by the plate to oppose the weight of the body. In clinical science and biomechanics, measuring ground reaction forces serves several purposes. Often, they are used with optical motion capture systems to generate a precise representation of the body, down to the millisecond and millimeter scale. A variety of scientific investigations hinge on this data—calculating torque at the knee, identifying pathological gait characteristics in a specific condition, or observing symmetry of gait while wearing a prosthesis. The precision of the measurements also makes them useful for creating and validating biomechanical models of movement. In that sense, lab-based systems are invaluable. However, researchers trade quantity and variety for precision—many force sensing systems rely on a “clean” foot strike to get good readings, and the types of movement possible in a sterile gait lab are not representative of the types of movement actually performed in daily life. Thus, instrumenting an insole with force sensors and wearing it during daily life would present the opportunity to observe qualitatively different, more realistic, data [20]. Several key populations can benefit from this. Individuals with diabetes, who often lack protective sensation and robust immune systems, are at increased risk of developing wounds on the foot. These wounds are at a high risk of infection, which can subsequently lead to hospitalization or even amputation, incurring massive costs along the way. Athletes put immense strain on their feet, and small deviations from optimal can place them at an increased risk of injury, deviations that can be detected using sensorized insoles. 77% of American adults have experienced foot pain [21] and half of those have reduced quality of life due to foot pain. We may conceive of several avenues by which force sensing wearable technology can address these widespread problems. Determining whether an individual’s foot and ankle biomechanics is pathological or simply idiosyncratic is incredibly difficult. For one person, excessive subtalar eversion may generate intolerable knee pain while for another it can be entirely asymptomatic. Forces on the foot may play a significant role, yet without the wide window of observation afforded by wearable devices, we may never understand the contribution that forces play. Lab-based measurement systems, like force plates, may be very precise, yet only can offer a sanitized version of human movement. Furthermore, wearable systems are cheaper and vastly more accessible. Smart analysis of force data may provide the user with early warning of injury risks and advise preventative measures—seeing an expert, wearing a certain type of shoe or insole, performing strengthening exercises, or avoiding certain activities. In the case of an individual with diabetes, force-sensitive insoles may have an important impact on health outcomes, for example, through early diagnosis, but a key technological hurdle remains before this can be realized at scale. Sensorized insoles today only sense forces normal to the foot, whereas the most destructive forces on the soft tissue of the foot are shear. Even minor shear forces repeated over time can damage the foot, though the threshold between normal and damaging is unknown due to the difficulty in measuring shear on the foot. Even in laboratory settings, measuring shear is difficult and requires specialized equipment. Sensorized insoles typically make use of force-sensitive resistors, which are cheap, low profile and easy to pack tightly together, and relatively durable considering the amount of force they are under. However, they only sense normal forces and require calibration to get precise force readings. A shear-sensing insole could be the first stage of defense against ulceration. Excessive forces over some window of time may be detected and the user warned, sparing them from a potential ulcer.

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Like with many other wearable technologies, sensorized insoles can generate orders of magnitude more data than lab-based systems, thus enabling new insights. However, there are still no widely-adopted commercially available options. Sensorized insoles on the market, such as Moticon [22], Tekscan, Pedar, and Medilogic [23], and Tactilus [24] are highly targeted to researchers, or perhaps elite runners. Like other wearable systems, they are limited by relatively low battery life and many rely on communication with a computer to extract recorded data. For researchers, these are acceptable costs, but for many consumers it remains to be seen what level of “hassle” is tolerable. Though lab-based systems will always be more precise and will retain their use, wearable systems are beginning to be available to translate that knowledge into practice in the real world [25].

14.3.2 Health and activity monitoring The traditional sense of health and activity lacks nuance and is ready to be disrupted by the widespread use of wearables. Health and activity are often assumed in the absence of an illness, but it is becoming increasingly clear that health and activity are multifaceted, highly personal, and poorly described by the binary concept of health or illness. The United States health care system is suffering from burdensome administrative bureaucracy that threatens the current practices and prevents expansion of services [26,27]. As a result, even the inadequate health monitoring standards of the past, such as an annual physical, are becoming increasingly inaccessible [28] and plagued by disparities in access [29]. Wearable devices provide individuals an opportunity to measure and understand their own health, and to improve our collective understanding of how signals from worn devices can improve health care, sports, and recreation. The way forward may have more to do with democratizing and distributing traditional health monitoring. Consider the example of the 2018 Scientific Report of the Physical Activity Guidelines Advisory Committee to US Health and Human Services [30]. It updated the previous recommendations from 2008 pertaining to “light-intensity physical activity” with duration less than 10 minutes. Previously, the report had lumped all forms of activity lasting less than 10 minutes, such as walking from a far-away parking space, into a baseline level of “background” movement undergone by people as a part of their regular lives. This seemingly trivial detail actually has enormous implications for public health, because the official recommendation of the government dismissed this sort of basic activity as a contributor to good health. For many Americans, the apparent uselessness of physical activity lasting less than 10 minutes created a barrier to beginning to improve their health by moving around more during normal activity. However, the 2018 report states “With the advent of devices to objectively measure physical activity of community-dwelling individuals during daily life activities in addition to exercise, it is becoming increasingly clear that light-intensity physical activity contributes to favorable health benefits, independent of those provided by moderate-to-vigorous physical activity.” It was the availability of ubiquitous wearable technology that made this understanding possible. Accelerometer-equipped devices like smartphones or smartwatches can count steps as a proxy for activity and are fairly ubiquitous. Further effort has been made to classify activity types based on inertial measurement unit (IMU) data, often using machine learning. Since the first wearable heart rate monitors became available in the 1980s, there has been a rapid and steady expansion of consumer-oriented health and activity monitors. The clip-on Fitbit classic (2009) marks the beginning of the current era, with a market size of $36 Billion in 2020 [31]. Beyond commercial impact, these devices provide an unprecedented opportunity for research, for instance to understand how people use these devices to change behavior, or monitor medical conditions [32]. Commonly available wearable activity trackers reasonably recognize common physical activities such as walking and running, with mixed success for more complex scenarios [33]. While these devices began primarily as step counters, they have now expanded into a wide variety of applications. The most common are estimating caloric expenditure, identifying more diverse classes of activities (running, sports, custom activities such as rowing, occupational activities [34], etc.), heart rate, and sleep [35]. Some commonly available devices also provide other functionality and measurements, including oxygen saturation, altitude, breathing, and skin temperature. Perhaps the most significant recent development in this device space is the introduction of electrocardiogram capabilities in wrist-worn activity trackers and smartwatches [36,37]. In addition to providing a more nuanced cardiac observation than simple heart rate, they present the first long-term monitoring available outside of the clinic. This means they can be used to detect atrial fibrillation that appears unpredictably and intermittently, and would be difficult to observe in a standard clinical visit [38,39]. The opportunity to detect diseases earlier has enormous potential impact on health outcomes [40]. These signals can be used for early detection of diabetes and stroke [41,42], and to estimate blood glucose levels from electrocardiogram [43]. What analogous measurements might be available using sensors in shoes? Achilles tendon thickening [44] or plantar tissue stiffness changes [45] associated with diabetes?

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14.3.3 Actuation and assistance The knowledgeable reader will note that this section is far from exhaustive. There is a diverse ecosystem of worn devices meant to assist the wearer by guiding the joints, providing power, resisting undesirable movements, and more. We aim, however, to orient the reader to the trends in orthotic aids, protheses, and more generally, wearable robots that provide active or semi-active actuation for movement [46]. There is considerable research interest in assistive devices intended to compensate for impairment or amputation [47 50], prevent injury [51,52], rehabilitate after injury or disease [53,54], and even for augmentation of unimpaired strength or dexterity [6,7]. Currently, devices are designed for specific narrow usage cases. Adding energy into the biomechanical system must be done with great attention to the neuroscience of movement and biomechanics, for example, reflexes, or spasticity dependence on forces. There are also considerable challenges with training. Since most active-power devices are not commercially available (but not all, e.g., [55]), it can be challenging to determine the usefulness of this class of hardware, and to disentangle the capabilities of the hardware with the influence of the control strategy. Actuation of the foot and ankle presents technical challenges because the gastrocnemius and soleus muscles produce a disproportionate share of total propulsive power in gait. Replacing these distal shank muscles with mechatronic elements presents daunting challenges in terms of weight and miniaturization. Active power robotic limbs or powered exoskeletons are intended to compensate for deficiency in that power output using motors, although they always also include passive elements, such as spring-like foot keels, to augment the power output for example, [56]. In more common clinical practice, passive foot-ankle orthoses are used to store power generated at the proximal joints and deliver it back to plantar push-off later in the stance phase of the gait cycle.

14.3.4 Haptics Haptics is the art of creating the experience of bodily interaction with the world. As a discipline, it is diverse and creative. For example, the feeling of interaction with objects is a fundamental aspect of industrial design. Creating the feel of a button, or the satisfying click when something closes securely, or the myriad tactile sensations arising from the textures, shapes, and movements of everyday objects, is a form of haptics [57]. But so too are devices that simulate interaction with virtual or teleoperated objects (e.g., [58 60]) or facilitate the kinesthetic experience of moving the body through virtual spaces [61,62]. Tactile confirmation of interactions with touchscreen user interfaces [63], the buzzing controllers of video games, wearables for rehabilitation [64], Braille language displays [65], and the feel of the steering wheel in modern cars [66] are all haptic applications. In this section, however, we will focus on the current state of purely wearable haptic devices.

14.3.4.1 Sensory substitution Worn haptics stimulates the sensory mechanoreceptors of the body for two general purposes. The first is sensory substitution, an effort to replace or augment sensory stimuli. Depending on the intended use case, this could range from full sensory replacement, to augmentation, to an additional modality for training or practice [67]. For example, when using a prosthetic limb, even an advanced robotic one, the sensory information of the physiological limb is unavailable. The gross forces developed by prosthesis interactions with the world appear at the socket interface to the residual limb in only an attenuated, diffuse way. We [68 71] and others [72 75] have addressed this by sensorizing the prosthesis and delivering a sensory substitution via worn haptics. There are even systems using implantable technologies to target nerves directly in lieu of stimulating the sensory mechanoreceptors, for example, [76,77]. Sensory substitution is critical for human computer interfaces as well. For example, worn virtual or augmented reality (VR, AR) systems typically provide rich visual stimuli, but lack the accompanying corporeal experience of body interaction. Recently there has been a renewed effort to create haptic interfaces to the hand: haptic gloves. For rehabilitation, hand exoskeleton gloves can be used to guide and assist people with spinal cord injury or other neuromotor injuries to perform movements, building strength and coordination, for example, [78 80]. One of the subtle, but important aspects of this strategy is that rehabilitation outcomes depend heavily on compliance—the regularity and duration of the performed exercises, and how engaged the user is in taking an active part in them. Integration of physical haptic aids with gamified rehabilitation protocols could facilitate adherence and engagement simply because they are fun, but also potentially because they richly engage, and therefore encourage plasticity of, the sensorimotor apparatus of the brain. Commercially available wearable haptic gloves are on the near horizon as well. These employ various simultaneous systems for tracking the movement of the hands and providing multisite stimuli to the skin [81]. They also use either tensile [82] or pneumatic compressive methods [81] to arrest the flexion of the fingers. This provides kinesthetic force feedback to muscle spindles and golgi tendon organs, the noncutaneous proprioceptive mechanoreceptors that elicit feelings of joint

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FIGURE 14.4 Haplets [86] are wearable, low-encumbrance wireless haptic stimulation devices that provide vibrotactile feedback for hand tracking applications in virtual and augmented reality.

movement and force generation—an oversimplification, see Refs. [83,84]. When coupled with the visual experience of interacting with objects, a powerful multisensory illusion of touch is created, increasing performance and qualitatively improving the experience. We have demonstrated that this illusion is robust to modest errors in correspondence between the apparent visual touch location and the felt tactile stimulation [85]. We have also developed more low-profile wearable finger haptic stimulation devices that can provide illusory touch during virtual object interactions [86]. Vibrotactile stimulations propagate from the dorsal fingernail, leaving the palmar surfaces of the hand free (Fig. 14.4).

14.3.4.2 Cueing and notification The other major use of wearable haptic stimulation is cueing: bringing the user’s attention to a piece of information, distilled to a useful cue instead of a complex tactile interaction [87]. By far the most common current application for wearable haptic cueing is in fitness trackers and smartwatches [88]. See Section 14.2.2 for a brief history and overview of what these devices typically measure. For haptic feedback, they generally, although not exclusively (See Ref. [89] for combined wristband squeeze and vibration) rely on vibrotactile cutaneous stimulation. They alert the user of health and activity events, such as the detection of a bout of exercise or milestones met such as a target number of steps in a day. Some devices are intended to be smartwatches, acting as an interface proxy to mobile phones. In contrast to activity trackers, they convey information from mobile phone alerts, act as a conduit to the phone with microphones, touchscreen interfaces, and even basic gesture interactions. We expect this trend to continue, with wrist-based bidirectional interfaces becoming an important interface to our worn computing environment [90 92]. Significant efforts have been made to develop wrist-worn aids for navigation and assistance of the visually impaired [93 97] (Fig. 14.5). These vary from systems to alert the user to navigational landmarks, to more ambitious efforts to sense the environment and convey useful information to the wearer. Similar cueing systems are useful for people experiencing impairment of vestibular sense [98 100]. We demonstrated that people wearing a cueing vibrotactile wristband could modulate their step length based on whether stimulation appears on the dorsal or ventral aspect of the wrist [101]. The idea is that upcoming gait events, such as trip hazards or terrain transitions, could be sensed by some other worn system. The movements of the user could be predicted and modulated by providing warnings or guidance to prevent trips or to serve as a tactile heads-up warning. This could become more important for augmented reality visual overlays and the concomitant demands on attention [102,103].

14.3.5 Motion capture Motion capture has become a foundational tool in human movement research but has historically required instrumented spaces and precise calibration using worn optical markers. The quality of these techniques is high, but the cost and preparation requirements have constrained the scope of observation. From the earliest use of cameras to answer biomechanical questions [104] to these room-scale systems, the ability to measure human movement has been miniaturizing and becoming more reliable. Today, we are beginning to have practical wearable versions of this technology in the

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FIGURE 14.5 Wrist-worn two-site haptic feedback system for indicating step length.

form of wearable IMU motion capture systems [105]. Wearable motion capture enables portability and makes data collection possible in more natural environments, such as outdoors or in the home. While this has thus far been primarily used in research, motion capture using wearable IMUs is beginning to be available for consumers directly [106], and if these systems fall into widespread use, it will dramatically transform the amount of human movement data that is available. Previously, most gait datasets available fell under the following categories: (1) Full-body kinematics and/or kinetics from level-ground (or treadmill) walking or running in in-lab environment comprising straight walking (10 m) different self-selected speeds [107 112], (2) staircase walking on in-lab staircases with 4 5 steps (see Ref. [20] for a summary of stair ambulation), and (3) datasets for human balance evaluation, for example, [113,114]. Recently there has been a flurry of datasets with more complex data, or with wider ranging activities in out-of-the-lab environments [20,115 119]. One notable study [115] collected data from 18 participants performing several activities with sensors attached at the thigh, shank, and feet. Another available dataset [116] includes full-body kinematics from different activities in a lab environment. Brantley et al. [117] captured full-body kinematics data on ramp, level-ground, and stairs, from 10 subjects. The ramp and 8-step staircase were constructed inside a gait analysis lab for that study, but another study [118] included data from a natural out-of the lab environment in three different terrain types. Recently, we have made an out-of-the-lab stair dataset available that includes 101 subjects descending an out-of-the-lab staircase [20]. Another study of note [119] collected more than 40 hours of unscripted daily life motion from 17 participants. The activities included a wide variety of activities, including: walking from one place to another, operating machinery, talking with others, manipulating objects, working at a desk, driving, eating, pushing and pulling carts, physical exercises such as jumping jacks, jogging, and pushups, sweeping, vacuuming, and emptying a dishwasher. The diversity of studies and lack of consistency with instrumentation, terrain, and activities raise a question about how many sensors are required to capture the movements of interest when pursuing this kind of work. The raw sensor measurements from worn IMUs require sophisticated models to translate into joint angles or segment positions. Thankfully, wearable systems enable capture of huge amounts of data with a fraction the work, making data-driven approaches viable, and indeed necessary, to make progress in the study and treatment of the foot and ankle. For example, we have shown [120 124] that ankle and knee kinematic trajectories can be reasonably predicted or estimated from the movement of the rest of the body or of the legs.

14.4

The future

Looking forward, there is great interest in surgical techniques (targeted reinnervation, regenerative peripheral nerve interface) that provide greater bidirectional informational exchange from user to robotic limb [125,126]. There is also continued progress in osseointegration, allowing prostheses to be grounded directly into the bone of the residual limb. This provides great opportunities to improve performance, but it certainly presents challenges in safety, risk of infection, and requires new surgical techniques [127]. Wearable technologies are in a period of vertiginous expansion. Historical data and recent trends in commercial products and research progress suggest a wave of maturing technologies that are making the leap from exotic to pedestrian. We expect daily life to include several forms of mechatronic and informational wearables within our lifetimes. Strap in, buckle up, and enjoy the walk.

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[96] Hong J, Pradhan A, Froehlich JE, Findlater L. Evaluating wrist-based haptic feedback for non-visual target finding and path tracing on a 2d surface. In: Proc. 19th international ACM SIGAC—CESS conference on computers and accessibility; 2017. p. 210 219. [97] Nair V, Olmschenk G, Seiple WH, Zhu Z. Assist: evaluating the usability and performance of an indoor navigation assistant for blind and visually impaired people. Assist Technol 2020;1 11. [98] Benini MJS, Bruinink M, Pekel AD, Talbott WA, Visser A, Markopoulos P. Restoring balance: replacing the vestibular sense with wearable vibrotactile feedback. Smart Healthcare Applications and Services: Developments and Practices. IGI Global; 2010. p. 283 301. Available from: http://doi.org/10.4018/978-1-60960-180-5.ch013. [99] Dozza M, Wall Iii C, Peterka RJ, Chiari L, Horak FB. Effects of practicing tandem gait with and without vibrotactile biofeedback in subjects with unilateral vestibular loss. J Vestib Res 2007;17(4):195 204. [100] Kingma H, Felipe L, Gerards M-C, Gerits P, Guinand N, Perez- Fornos A, et al. Vibrotactile feedback improves balance and mobility in patients with severe bilateral vestibular loss. J Neurol 2019;266(1):19 26. [101] Sie A, Fisher C, Karrenbach M, Caraballo C, Case E, Muir B, et al. Timing of haptic cues for stride adjustment in mobility task. sieWristHaptics 2022. In preparation. [102] Bosman S, Groenendaal B, Findlater JW, Visser T, de Graaf M, Markopoulos P. GentleGuide: an exploration of haptic output for indoors pedestrian guidance; 2003. p. 358 362. Available from: https://doi.org/10.1007/978-3-540-45233-128. ,http://link.springer.com/10.1007/9783-540-45233-128.. [103] Karuei I, Maclean KE. Susceptibility to periodic vibrotactile guidance of human cadence. In IEEE haptics symposium HAPTICS, IEEE Computer Society; 2014. p. 141 146. Available from: http://doi.org/10.1109/HAPTICS.2014.6775446. [104] Muybridge J. The horse in motion. Nature 1882;25(652) 605 605. [105] Schepers M, Giuberti M, Bellusci G, et al. Xsens mvn: consistent tracking of human motion using inertial sensing. Xsens Technol 2018;1(8). [106] Notch wearable motion capture, ,https://wearnotch.com/. [accessed 1.3.22]. [107] Moore JK, Hnat SK, van den Bogert AJ. An elaborate data set on human gait and the effect of mechanical perturbations. PeerJ 2015;3:e918. [108] Fukuchi RK, Fukuchi CA, Duarte M. A public dataset of running biomechanics and the effects of running speed on lower extremity kinematics and kinetics. PeerJ 2017;5:e3298. [109] Fukuchi CA, Fukuchi RK, Duarte M. A public dataset of overground and treadmill walking kinematics and kinetics in healthy individuals. PeerJ 2018;6:e4640. [110] Burdack J, Horst F, Giesselbach S, Hassan I, Daffner S, Sch€ollhorn W. A public dataset of overground walking kinetics in healthy adult individuals on different sessions within one day. Mendeley Data 2020;1(2). [111] Schreiber C, Moissenet F. 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Chapter 15

Integrated Laboratories for Pursuing Pedal Pathologies Oliver Morgan1, Rajshree Hillstrom2, Jinsup Song3, Robert Turner4, Marian T. Hannan5, Yvonne M. Golightly6, Scott J. Ellis4, Jonathan Deland7 and Howard J. Hillstrom4 1

Stryker, Centennial Park, Elstree, United Kingdom, 2Biomed Consulting, New York, NY, United States, 3Temple University School of Podiatric

Medicince, Philadelphia, PA, United States, 4Hospital for Special Surgery (HSS), New York, NY, United States, 5Harvard Medical School, Hebrew SeniorLife, Boston, MA, United States, 6University of North Carolina, Chapel Hill, NC, United States, 7Orthopaedic Surgery, Weill Cornell Medical College, Attending, Department of Orthopedics - Foot and Ankle, HSS, New York, NY, United States

Abstract The study of complex human movement pathologies requires a multidisciplinary comprehensive approach that considers both normal and pathologic structure and function. Common areas of clinical research include epidemiology, in vivo experimentation, in vitro experimentation, and in silico simulation. Historically, these have been independent on one another; however, integrating these areas is recognized as a strength. Each may provide novel information and support the other to advance our understanding of pathology, which may be particularly important for under-recognized and understudied disorders associated with the foot and ankle. In the context of pedal pathologies, this chapter reviews the general concepts and specific examples in the areas of epidemiology, in vivo experimentation, in vitro experimentation, and in silico simulation. Furthermore, a case study in osteoarthritis of the first metatarsophalangeal joint (hallux rigidus) is presented to demonstrate integration of these different areas to inform and advance clinical strategy. This approach of integrating laboratories is useful for providing mechanistic insight into the onset and progression of neuromusculoskeletal disease. Such insight can spawn future studies that focus on improved treatment strategies, tools to personalize treatment planning, and ultimately prevent disease or injury.

15.1

Introduction

Investigation of pathologies has become a multi-discipline endeavor. Feet, in particular, are often impacted by structure (e.g., planus, normal, or cavus) and function (e.g., excessive eversion, normal, or excessive inversion). These biomechanical variables may be assessed in vivo or with large data sets, using epidemiological approaches. A more intrinsic perspective may be obtained from cadaveric studies or using computational modeling. To determine the prevalence of a certain pathology or incidence of disease prospectively (e.g., hallux rigidus or hallux valgus), then an epidemiological study is required. However, to determine whether individuals with first metatarsophalangeal (MTP) joint osteoarthritis are overloading their hallux, an in vivo pedobarographic analysis is required. Moreover, to understand the contact mechanics of the first MTP joint, a cadaveric or a computational model is required. Pedal pathologies are among the least understood and understudied in the human musculoskeletal (MSK) system. While significant research focus has been given to diseases of the hand, hip, and knee, comparatively modest attention has been directed toward the foot. Advances in interdisciplinary collaborations over the past decade have led to a more streamlined approach in the pursuit of understanding foot problems. Our approach to research of pedal pathologies, which integrates the various knowledge bases of engineers, scientists, and clinicians will be discussed in this chapter.

Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00054-8 © 2023 Elsevier Inc. All rights reserved.

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Most investigational teams in modern medicine and science are multidisciplinary, recognized as the best way to achieve clinically relevant results from research. Disciplines including epidemiology and biomechanics, which historically have been independent from one another, are beginning to recognize strength in integration. Epidemiologists can identify trends in disease and high-risk populations, but their methodologies can be limited by case definitions and the availability of appropriate equipment [1]. Contributions from biomedical engineers can strengthen the quality of epidemiological outputs by improving the design of precision tools. Novel technologies are the foundation of research, and their development requires creativity and in-depth understanding of the problem at hand. Integration of different research fields is beneficial to advance clinical strategy and inform innovative treatments for pedal pathologies. Thus, a system has begun to form, where epidemiologists advise on clinical need, biomedical engineers build equipment for more precise measurements, and clinicians translate their findings for patient intervention. While this is a simplified explanation of integrating laboratories for foot and ankle care, it conveys the basic framework necessary to pursue pedal pathologies.

15.2

Our method of approach

Advances in interdisciplinary collaborations over the past decade have led to more streamlined methodologies for understanding foot problems. Our approach to research, which integrates the various knowledge bases of engineers, scientists, and clinicians will be discussed in this chapter. The approach recommended by our investigational team is to integrate the following laboratories and expertise toward the pursuit of foot and ankle pathology (Fig. 15.1). Data from each laboratory can inform the others in terms of biomechanical and clinical outcomes, experiment design, disease and treatment mechanisms, predictive computational model development and validation, and translation to clinical care. Our approach consists of the following components: 1. Epidemiology: Study of population prevalence and incidence trends in health. 2. In Vivo Experimentation: Determine the physiological structure-function relationships of the neuromusculoskeletal system. 3. In Vitro Experimentation: Define the mechanical properties of biological tissues and measure their internal mechanical behaviors. 4. In Silico Simulation: Perform predictive analyses for pathologic function and pre-operative planning.

FIGURE 15.1 Graphical representation of the integrated laboratories methodology in clockwise order. (Top left) Epidemiology: study of foot type among United States Military Academy (USMA) cadets; (Top right) In vivo: gait analysis and ground reaction force measurement; (Bottom right) In vitro: measurements of the medial metatarsosesamoid ligament in a cadaveric specimen; (Bottom left) In silico: resultant stress in the structures of the first ray from simulated physiological loading.

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Integrated laboratories

15.3.1 Epidemiology Epidemiological studies have been used to determine the prevalence of foot structure, function, flexibility, predict injury incidence, and observe longitudinal changes in disease. The design of such research can involve short-term randomized controlled trials, mid-term retrospective analyses, or long-term population-based studies. The usefulness of these methods depends on the study design. For example, a randomized controlled trial may be useful to investigate a new interventional drug, while long-term population-based studies could help determine trends in disease and identify groups most at risk. Foot problems are common, and these rise with increasing life expectancy. The prevalence of pedal pathologies can vary greatly between different studies; however, the figure appears to be between 71% and 87% among older adults [2 4].

15.3.2 In vivo experimentation In vivo experiments encompass a wide variety of methodologies, and all of which pertain to the study of living subjects. Like many of the methods detailed in this chapter, the experimental design of in vivo research depends on the point of interest. The most common methods for understanding pedal pathologies are gait analysis and plantar pressure assessments. Human movement is an important consideration in understanding neuromuscular interactions and kinematics of the skeletal system, especially when designing treatments for injury or disease. In vivo experiments to measure gait range from motion capture, force plates, and electromyography.

15.3.2.1 Gait analysis Gait analysis techniques will typically use surface-mounted markers to track human motion. In the foot, rigid segment models have been utilized, reducing the overall complexity of its structure [5 7]. Nester et al. identified two fundamental sources of error: (1) simplification of foot anatomy and (2) skin movement artifact [8]. To understand how the assumption of the foot to be a rigid body affected its kinematic description, they compared the results from a traditional skin-mounted marker system to bone anchored markers. Although screwing pins to the bones of the foot is not a viable technique in a clinical setting, this method provides knowledge of the specific limitations associated with skin-mounted markers. This is particularly important for the foot due to the complex interactions of its kinetic chain during gait. Unlike the knee, which acts like a rolling contact hinge, or the hip, which acts like a ball-and-socket, motion in the foot can vary depending on each specific joint. The simplification of its structure can have implications on its representation in in vitro and in silico studies.

15.3.2.2 Plantar pressure assessments Plantar pressure can be used to describe foot function (Fig. 15.2). Planus feet have elevated peak plantar pressures and pressure time integrals beneath the medial arch, central forefoot, and hallux, while cavus feet have higher pressures at the heel and lateral forefoot [9]. The center of pressure (COP) excursion index (CPEI) can be used to quantify function as well [10]; planus feet have a smaller CPEI value and a less concave COP curve, while cavus feet have a larger CPEI value and a more concave COP curve [11].

15.3.2.3 Measures of foot structure There have been numerous measures of foot structure proposed and implemented, including the foot posture index, which is a measure of arch structure based on six visual observations [12]. Other measures include the navicular height and the navicular drop [13], which quantifies the relationship between the navicular tuberosity and the ground, as well as the arch index, a foot print-based ratio related to the amount of ground contact [14].

15.3.2.4 Other measures Other in vivo analysis techniques include radiographic assessments of foot pathology (i.e., 2D static alignment) [15,16], weight-bearing computed tomography (CT) scans (3D static alignment) [17,18], and biplanar fluoroscopy (dynamic bone motion) [19,20]. Electromyographic activity of the lower limb extrinsic and intrinsic musculature is also useful for understanding foot function [21,22].

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FIGURE 15.2 Plantar pressure measurements.

15.3.3 In vitro experimentation In vitro testing is necessary to explore aspects of the human MSK system, which would otherwise be unethical or too dangerous to perform in vivo.

15.3.3.1 Histology Histology is the study of microscopic tissue structures and can be used to define the material properties of bone, cartilage, ligament, tendon, etc. Athanasiou et al. and Liu et al. are some of the few researchers to investigate cartilage properties in the foot and ankle [23 25]. Biphasic creep indentation experiments are used to determine the material properties of cartilage, using cadaveric specimens. To ensure all specimens were healthy, both investigators used the India ink staining technique. The creep recovery responses of cartilage specific sites were measured using an automated creep indentation apparatus. The tissue material properties, in these studies, were described by linear biphasic theory, accounting for aggregate modulus (HA in MPa), which is the tissues compressive modulus, the Poisson’s ratio (dimensionless, more frequently denoted as: v), the shear modulus (µs in MPa), and permeability (k in 10 15 m4/N.s), which indicates the interstitial fluid pressurization. Cross sections of articular cartilage were also obtained to allow for analysis of chondrocyte arrangement through the different zones.

15.3.3.2 Cadaveric simulators Cadaveric gait simulators are designed to reproduce foot kinetics and kinematics in vitro. Depending on the question to be addressed, some designs may be simplistic mechanical jigs (i.e., for static loading), while others may attempt to recreate the complex functions of the foot using electronic-actuated robots (i.e., for dynamic loading). In vitro simulators offer descriptions of foot and ankle kinematics where estimates from gait analysis may be incomplete due to the assumption of rigid body kinematics. Moreover, they enable the researcher to recreate the foot’s loading environment and investigate healthy, pathologic, and surgically treated articular contact mechanics. Ultimately, their uses revolve around repeatable and controllable simulations of gait and measurements of internal biomechanics. Static simulators are useful as feasibility tools, to understand the interactions of different soft-tissues in a more controlled environment than dynamic simulators. Kim et al. used a custom-made static simulator (Fig. 15.3) to investigate the effect of a 3-mm Moberg osteotomy on first MTP joint contact mechanics [26]. By manipulating the passive effect of the windlass mechanism on hallucial dorsiflexion, the researchers were able to investigate the arc of first MTP joint motion, pre- and post-3-mm Moberg osteotomy, and the commensurate changes in contact pressures, contact areas, and COP locations. Early dynamic simulators tended to simplify the function of the foot and ankle by reducing the number of tendons under actuation, constraining the tibia to a sagittal rotation, and scaling velocity and loads [27 29]. Although able to

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FIGURE 15.3 (A) Experimental custom-made test jig for the first ray with specimen fixed in position at the metatarsal base; (B) insertion of a TekScan K-Scan 6900 pressure sensor (TekScan Inc, Boston, USA) into the first metatarsophalangeal joint cavity; and (C) visualization of the resultant joint contact pressure pattern and magnitude.

provide more detailed information than static simulators, these dynamic conditions were also limited to the study of certain aspects of foot and ankle biomechanics. Aubin et al. argued that, although important, the requirement of early gait simulators to perform a large number of cadaveric simulations or to generate a broad dataset was not well suited to the short lifespan of cadaveric tissues [30]. They introduced a dynamic simulator which employed fuzzy logic ground reaction force (GRF) control and operated at a realistic velocity. This simulator [31 33], or one similar to it [34,35], has been used to study a range of subjects, including orthopedic surgeries.

15.3.4 In silico simulation The term in silico was first introduced in the late 1990s, alluding to Latin phraseology in science: in vivo and in vitro [36]. This pseudo-Latin phrase is defined as “(of scientific experiments or research) conducted or produced by means of computer modeling or computer simulation.” In silico research includes finite element (FE) modeling and MSK modeling. Both platforms simulate virtual environments to predict human kinetics and kinematics. Such simulations are developed from reconstructed radiological image data, e.g., CT or magnetic resonance imaging (MRI). These models may be generated from the anatomy of cadaveric specimens, living subjects, or generic scaling algorithms. FE models are used to represent the foot’s anatomical geometries, predicting soft-tissue stress and strain by simulating physiological loading. It has been used to study bone and joint biomechanics, surgical interventions, and implant design. A few decades ago, commercially available software platforms, such as Abaqus (Dassault Syste`mes, Ve´lizyVillacoublay, France) and Ansys (Canonsburg, Pennsylvania, USA) were typically used to develop FE models and run

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analyses to predict contact mechanics in bones and joints. More recently, FEBio (University of Utah, Utah, USA), has introduced through a collaborative development from the Musculoskeletal Research Laboratory at University of Utah and the Musculoskeletal Biomechanics Laboratory at Columbia University, to provide researchers with a publicly available open-source platform to conduct FE investigations. Nonetheless, the requirement of assumed boundary/loading conditions is a limitation of any FE software, and, hence, a researcher’s ability to predict contact mechanics in the tissues. This may result from constrained degrees of freedom or underestimated contact forces measured from inverse dynamics. Independently, MSK modeling was introduced to estimate neuromuscular control, simulating muscle-driven multibody dynamics. OpenSim and Simbody (University of Stanford, Stanford, USA) enable non-invasive estimates of muscle tension and joint contact forces. The software, as well as various projects are accessible through Simtk (Simtk.org), an open-source project-hosting platform for in silico investigators. More recently, FE and MSK modeling have been combined, providing more accurate predictive capabilities, whereby joint contact forces during static or dynamic activities may be used to drive the FE model to improve the estimates of stress, strain, and contact mechanics in joints and other tissues.

15.3.4.1 Medical image processing To build anatomical-based FE or MSK models, 3D tissue geometries are developed from radiographic images (Fig. 15.4), such as CT or MRI datasets. Several platforms have been developed to support 3D image processing and model generation. Mimics (Materialise NV, Leuven, Belgium) and Simpleware (Synopsys, Exeter, UK) are two commercial packages widely used in research and industry. Both offer complete segmentation and model generation tools, starting from importing the raw datasets through to geometry refinement and FE meshing. Several other freely available software packages can be used, depending on the user’s requirements. Image segmentation is usually performed semi-automatically. Although some automated segmentation tools have been developed [37], their algorithms are reliant on image quality and vary depending on anatomical site. Manual segmentation is a particularly time-consuming stage of the modeling process. The geometry of bone can be considered easy to segment thanks to its large surface area and, particularly in CT images, clear contrast and conversion of Hounsfield values to gray scale, which helps with automation of the segmentation process. Smaller geometries, such as cartilage and ligaments, are more difficult due to low contrast with the surrounding structures. Inaccurate segmentation of articular surfaces and ligaments may result in erroneous predictions of joint contact mechanics. Segmenting the foot’s anatomy can be particularly challenging due to the small size of its individual soft tissues. In some cases, MRI-image resolutions, coupled with the small size of ligaments in the forefoot, can mean that accurate

FIGURE 15.4 Segmented anatomical geometries of the great toe (hallux) from MRI imaging.

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segmentation of these structures are not feasible. It may therefore be necessary for a researcher to idealize their functional contributions later in the modeling process.

15.3.4.2 Finite element modeling FE modeling is a numerical method to predict the performance of a structural system, such as a reconstructed joint. To increase the accuracy of an FE model prediction, the geometrical representation of the anatomical tissues must be accurate, the interaction between the tissues and physiological loading conditions must be realistic, and the element size must be small enough. The rise of powerful computers and commercially available FE software in the early 1980s provided biomedical researchers a new tool for exploration, bridging the gap between in vivo and in vitro experiments [38]. The primary difficulty of studying tissue mechanics in vivo are the potential risks to and therefore ethical considerations of using human subjects. These issues can be circumvented with FE modeling, which can predict detailed biomechanical information, including internal tissue stress and strain (Fig. 15.5). Moreover, if we take the example of patient-specific surgical optimization, one FE model can be used to simulate and assess the efficacy of multiple surgeries, enabling the researcher to non-invasively compare surgical outcomes from the biomechanical perspective. Nakamura et al. investigated pedal mechanics using an FE model to understand joint stress in the foot under weightbearing load [38]. This simplistic model simulated a variety of shoe conditions during mid-stance. It was not until much later, in 2005, that Cheung et al. published a 3D FE model of the entire foot-ankle complex [39]. A series of investigations were conducted using this model, spanning analyses of soft-tissue properties [40], basic pedal mechanics [41], and plantar fascia surgery [42]. These early examples demonstrated the benefits of in silico modeling in supplementing traditional methods of research. Many FE models represent the entire foot; however, simulating the physiological conditions of all 33 joints in the foot can be computationally demanding and can diminish the practicality of predicting surgical outcomes, when analyzing specific structures of the foot. Reconstruction of smaller anatomical segments of the foot can be necessary to allow detailed physiological representations and interactions of the different tissues, and enhance clinical knowledge, while reducing the temporal-cost of development and analysis [43].

FIGURE 15.5 Example of finite element model of the great toe (first metatarsal, proximal, and distal phalanx) under simulated weightbearing load. (A) Experimental set up; (B) segmentation from MR images; (C) 3D reconstructions of anatomy; (D) meshed FE model.

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15.3.4.3 Musculoskeletal modeling Musculoskeletal models comprise anatomical sections which are connected by joints with prescribed degrees of freedom. Insertions and pathways of muscles and ligaments are modeled, where external loads from gait analysis are input to determine their resultant forces [44,45]. Musculoskeletal models of the lower limb have tended to focus on the hip and knee joints [45 47]. Arnold et al. published a lower limb MSK model, which included the phalanges, metatarsals, calcaneus, and talus of the foot, with corresponding coordinate systems [48]. The model was used to investigate variation in muscle forces and joint moments over a variety of body positions. Of the MSK foot-specific models that have been developed, Saraswat et al. presented a particularly detailed model which included the intrinsic muscles and ligaments of the foot [49]. However, this model remained a generic representation. Prinold et al. developed a patient-specific method for the estimation of ankle joint forces in patients with juvenile idiopathic arthritis [50], according to the framework of the modified Oxford Foot Model [51]. This provided the foot and ankle with seven degrees of freedom: three at the ankle, three at the hindfoot and forefoot, and one between the forefoot and toes. Kim et al. investigated joint moments and contact forces in the foot during walking using an AnyBody MSK model [52]. Oosterwaal et al., developed a more complicated, 26-segment kinematic model of the foot to gain insight into intrinsic foot kinematics [53]. The model, known as the Glasgow-Maastricht foot model, is one of the most detailed MSK models available for the foot. Distinct motion patterns were found for each joint.

15.3.4.4 Sensitivity studies Sensitivity analyses aim to reduce the errors within an in silico model. This may be for 3D reconstruction, material properties, or mesh size. Mesh sensitivity involves a convergence study, usually performed to test the accuracy-time benefit of a simulation by determining the least change in results compared to the fastest simulation time. For those interested in contact mechanics, the mesh of articular cartilage may be assigned a fine density, while that of the bone is increased to reduce the total time in the FE solver. Additionally, the mesh biasing toward the articular and chondral surfaces can be performed, reducing the total number of elements in an FE model while retaining focal areas of dense mesh. More commonly, sensitivity analyses of material properties and mesh density are presented [43,54 56]. Optimization of material properties with measured data can supplement variations in material characteristics between subjects and specimens, compared with mean values usually presented in histological research. The resolution of a radiographic image can also be a source of error. If the voxel dimensions of the image are large it can result in deviations from the true geometry of the subject or specimen. Furthermore, manual segmentation of a 3D image can also lead to errors due to the researcher’s ability to identify different soft tissues [57,58].

15.3.4.5 Validation When considering the applicability of any model, it is important to remember that they are predictive tools that require assessment of validity and reliability [43,55,56]. The main challenge of in silico research is the uncertainty relating to material properties and macro-geometry of biological tissues. Validation of their outputs are normally used; ensuring the result is correctly interpreted before its predictions can be considered clinically valuable [59]. The process of in vivo validation can be considered direct or indirect. A direct validation collects data from the same subject used to build the model [43]. An indirect validation uses gait data from the literature, matching their results with deviations from mean measurements [55,60]. FE investigations of the foot have generally involved healthy humans. This can be attributed to its interface with the ground, and hence, the ability to compare measured to predicted plantar pressures. Measurements of plantar pressure may usually be taken during gait analysis by a dynamic in-sole system or pressure plate [54,59]. This validation method is limited by the assumption that the predictions of kinematics, and joint pressures are accurate, without validating for soft-tissue geometries or material properties [43,56]. One example of these different validation methods being utilized was demonstrated by Morales-Orcajo et al. who estimated material parameters for tendons in the foot based on data collected from 100 uniaxial tensile tests [56]. Vertical displacement of the foot’s arch was then indirectly compared with experimental measurements from the literature. Subsequently, the authors further validated their model using direct in vivo plantar pressure measurements. In vitro validation is a more complicated process than that for in vivo studies. It may require integration of multiple expertise including design engineering/knowledge of cadaveric simulators, software programming, anatomical dissection, and all in addition to FEA. The most common in vitro validation metric is intracapsular pressure sensor measurements [43,60] though some investigators may use a cadaver-based kinematic marker system [35].

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Wong et al. employed an indirect contact validation in their study of metatarsocuneiform (MTC) joint arthrodesis [60]; however, in instances that an in vitro validation is performed, a direct validation is always advised. This is because of joint-specific morphology, selected loading parameters, foot alignment, and specimen-specific material properties. In vitro validation studies are complicated, but validated models can provide important clinical recommendations. Some in vitro validation procedures may define the in situ ligament or bone strain as the parameter of interest. Fung et al. presented an experimental validation of FE predicted metatarsal strain [61]. A custom-made aluminum fixture was designed to position the metatarsals at various angles in an MTS materials testing machine (MiniBionix 858, MTS Systems, MN, USA). Corresponding FE models were developed and the measured versus predicted bone strains were compared to establish model validation.

15.4

Case study of the integrated laboratories concept to the study of hallux rigidus

The following descriptions summarize the dissertation of Oliver Morgan, PhD which applied the integrated laboratories concept to the study of first ray hypermobility and first MTP joint osteoarthritis (or hallux rigidus as it is commonly known) [43,62 64]. It has been theorized that the planus foot type and first ray hypermobility may be interrelated mechanisms, initiating hallux rigidus in the early stage [11,65,66]. The research hypothesized that during gait, the first metatarsal of a hypermobile planus foot will translate excessively in the superior direction, causing the foot to evert, and redistributing the body’s weight. Once the first ray reaches maximum elevation, the medial band of the plantar fascia may become maximally taught, causing the first MTP joint to undergo increased loading and dorsal articular impingement. The overarching theory considered that planus foot type and first ray hypermobility could permit repetitive excessive loading to the first MTP joint and eventual degenerative changes indicative of hallux rigidus. We present an overview of what was learned from each of the four integrated laboratories.

15.4.1 Epidemiology Despite the frequency of foot injuries and diseases [2 4], a systematic review of epidemiology studies for foot osteoarthritis found just 27 articles published before 2010. A similar systematic review, published two years earlier, found 176 epidemiology studies for hand OA and 190 for the knee [67]. Compared to the hip, knee, and hand, little population data exists for osteoarthritis of the foot. We addressed this issue by conducting a study of osteoarthritis of these anatomical joints using population data (http://www.webarchives.nationalarchives.gov.uk) from the NHS Hospital Episode Statistics in England [62]. During the years of 2000 to 2018 we included 3,143,928 individuals with osteoarthritis of the first MTP, first carpometacarpal (CMC), hip and knee joints (Table 15.1). Population statistics were computed for joint specific osteoarthritis stratified by age and sex. Women had almost double the prevalence of osteoarthritis at the first MTP, CMC, hip, and knee joints compared with men. The incidence of osteoarthritis increased significantly at the first MTP, first CMC, hip, and knee joints per year from 2000 to 2018 (Table 15.2). Although the first CMC joint had the highest incidence of disease, the first MTP joint had comparable incidence to the hip and higher incidence than the knee. The prevalence of osteoarthritis at the first MTP joint substantially increased from 2000 to 2018 (Fig. 15.6). More interesting though was that this group exhibited a bimodal distribution suggesting that younger and older populations were involved. These findings underscore the importance of tracking the epidemiology of first MTP joint osteoarthritis and hallux rigidus.

TABLE 15.1 Prevalence of osteoarthritis at the first MTP, first CMC, hip, and knee joints, 2000 18. Prevalence

Joint First MTP

First CMC

Hip

Knee

60,986

88,178

1,222,446

1,772,318

Men

32%

24%

39%

43%

Women

68%

76%

61%

57%

All (N)

Note: First MTP osteoarthritis presented comparable diagnosis numbers to the first CMC joint.

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TABLE 15.2 Incidence of osteoarthritis at the first MTP, first CMC, Hip, and Knee joints, 2000 18. Incidence

Joint First MTP

First CMC

Hip

Knee

3.8% (3.0, 4.6)*

10.9% (10.1, 11.7)*

3.8% (2.9, 4.7)*

2.9% (2.2, 3.6)*

Men

3.6% (2.8, 4.5)*

11.9% (10.7, 13.2)*

4.3% (3.2, 5.5)*

2.7% (2.0, 3.4)*

Women

3.9% (3.0, 4.7)*

10.7% (9.8, 11.7)*

3.8% (2.9, 4.7)*

3.1% (2.3, 3.8)*

All

Note: Significant change in incidence for P , .05 indicated by *. In the regression analysis, AAPC (%) and 95% CI in crude rates per 100,000 population were used.

FIGURE 15.6 Prevalence and age distribution of first metatarsophalangeal joint osteoarthritis from 2013 18.

15.4.2 In vivo experimentation Pbes planus is a known risk factor for the development of hallux rigidus or osteoarthritis of the first MTP joint [68]. Several investigators have observed greater odds of developing stress fractures during military training in subjects with pes planus [69,70] or pes cavus [71]. Foot type has also been correlated with decreased joint range of motion, flexibility [72,73], and increased hallucial loading [11]. These mechanical factors affecting people with pes planus (i.e., decreased joint range of motion, flexibility, and increased hallucial loading) likely contribute to higher first MTP joint stress [11]. Repetitive excessive loading of the joint may lead to cleavage of cartilage matrix and at the end-stage, erosion of the cartilage overlaying subchondral bone [74,75]. Previous research found excessive hindfoot eversion and increased medial column loading of the planus foot [11,73,76 78]. In particular, the peak pressures beneath the first MTP joint are decreased, while those beneath the second MTP joint are increased in planus compared to well aligned feet [11]. Specifically in the planus foot, this suggests that hypermobility of the first ray may be an important factor for the development of several pedal pathologies. In our original study of foot type biomechanics in asymptomatic individuals [11], dorsal displacement of the first ray (mobility) was not measured (Fig. 15.7). There were no commercially available devices to make measurements of first ray mobility that had withstood the rigors of intra- and inter-rater reliability. Morgan et al. designed and prototyped a device to measure mobility and position of the first ray (MAP1st). The rationale for developing this device was that foot type and first ray hypermobility may interact in the development of first MTP joint osteoarthritis. The basis for this theory has existed both, in earlier research [11] and in our earlier epidemiological finding of a biomodal peak in younger patients who exhibited hallux rigidus. We hypothesized an additional risk profile for hallux rigidus, other than wear and tear of the articular soft tissues in old age, which may be biomechanical in origin. One plausible theory for developing osteoarthritis of the first MTP joint is that individuals with a planus foot type and first ray hypermobility could lead

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FIGURE 15.7 First ray mobility where each white circle denotes the first through fifth metatarsal heads. A change from static forefoot alignment, which illustrates the equilibrium position for the first ray, to first ray mobility that occurs in response to first MTP joint plantar loading is shown.

FIGURE 15.8 MAP1st (left and right) prepared for testing, with each of the subject’s feet positioned and clamped in the device. The Arduino microcontrollers of each device are plugged into a laptop to interface the custom-written code for upload and testing. The program was written to apply cyclic loads of 25 N to control for the recent strain history of the first ray soft tissues.

to a tautening of the plantar fascia under load bearing. The contact forces within the first MTP joint in turn may promote an excessive stress in the dorsal aspect of the joint that damages cartilage matrix and the corresponding tissues that comprise the diarthrodial joint. MAP1st (Fig. 15.8) was designed as a computer-controlled device that could apply precise loading (e.g., 50 N) beneath the first MTP joint in the superior direction (PCT/US21/22791, 2021). The second MTP joint was mechanically grounded so that the movement of the first ray was isolated. The graticule permitted the measurement of first ray mobility between the loaded and unloaded conditions. The actuator was driven by custom code running on an Arduino microprocessor. To assess the reliability of MAP1st, 25 participants (50 feet) were recruited and tested by two different raters. Test retest and remove replace conditions were assessed. MAP1st mobility was measured with and without normalization to foot length. We also included partial weight bearing (PWB, seated) and weight bearing (WB, standing) conditions as well as resting calcaneal stance position (RCSP) and subtalar joint neutral positions. Intraclass correlation coefficients (ICC(2,1)) were computed for intra-rater and inter-rater reliability determination along with Bland-Altman

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FIGURE 15.9 The arch index system used to assess and stratify individuals based on planus, normal, or cavus foot type.

plots [64]. MAP1st intra-rater ICC values were $ 0.85. Inter-rater reliability was 0.58 for normalized first ray mobility assessments in either PWB or full WB and in RCSP. Reliability of a hand-held ruler device was also performed as a clinical comparator which was similar for intra-rater reliability (ICC $ 0.85) but unacceptable for inter-rater reliability (ICC 5 0.06). MAP1st provided reliable measurements of first ray mobility in PWB and WB in RCSP. This tool should help explore relationships between foot type and first ray mobility in asymptomatic and pathologic feet. Foot type is associated with many pedal pathologies. Both hallux valgus and hallux rigidus are associated with pes planus [68,73,79,80] while hammertoes have been associated with pes cavus, and injuries with both pes planus and cavus foot types [69]. The role of hypermobility of the first ray (first metatarsal and medial cuneiform) is less clear although often anecdotally described as playing a role in the development of hallux rigidus and hallux valgus. To address the question “Is the planus foot type associated with first ray hypermobility?,” twenty asymptomatic participants with planus (n 5 23) and normal (n 5 17) feet were recruited [63]. Static foot structure (arch height index (Fig. 15.9), arch height flexibility, first MTP joint flexibility, and first ray mobility) was measured for each participant’s feet. These same subjects were also stratified by hypermobile first rays (n 5 14) and non-hypermobile first rays (n 5 26). First ray hypermobility was defined as when first ray mobility $ 8 mm in superior translation. In this cohort, 86% of participants with hypermobile first rays were classified as having pes planus feet according to their arch height index (AHI , 0.345). There were significant differences in arch height flexibility and weight-bearing first ray mobility. Weight-bearing first ray mobility and first MTP joint laxity were associated with PWB first ray mobility (R2 5 0.38). Furthermore, the planus foot type was found to be associated with first ray hypermobility. The interaction between flexibility of the first MTP joint, first ray mobility, and foot type makes sense when considering the windlass mechanism since the plantar fascia is put into tension when vertical load is applied to the foot. In feet with normal first ray mobility, flexibility of the first MTP joint during dorsiflexion is decreased. When the first ray is hypermobile the medial arch lowers into greater pes planus which further decreases first MTP joint flexibility. This effect has been theorized to impose increased stress upon the dorsal aspect of the first MTP joint [54,55]. The measures of foot structure as applied in this study may help us understand the sequela associated with symptomatic pathologies of the foot.

15.4.3 In vitro experimentation FE modeling of the foot may be used to predict and visualize internal soft tissue stress and strain distributions or perform parametric studies. However, the perceived reliability of such models should always be governed by verification and validation against experimental data. Physical testing can be considered essential in the evaluation process for reliable in silico modeling. Therefore, to predict the theoretical influence of first ray hypermobility on first MTP joint loading, our investigational team designed a custom force-controlled cadaveric test-rig for physiological loading of the medial forefoot and intracapsular TekScan (TekScan Inc, Boston, MA, USA) pressure sensor measurements of first MTP and first MTC joint contact mechanics (Fig. 15.10).

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FIGURE 15.10 (A) Isometric view of the test-rig with the cadaveric foot ready for testing. The medial, middle, and lateral cuneiforms are fixed to a pivot mechanism at the anterior portion of the device which allows sagittal rotation, while the calcaneus has been fixed to a posterior component, with all DOF constrained; (B) measurements of the metatarsal angle using a goniometer to define the arch-alignments of the normal and planus foot types; (C) A SolidWorks illustration of the initial design concept for the cadaveric medial forefoot test-rig, shown from a lateral cross-sectional view. The forces acting on the system are denoted using red arrows while rotating components are denoted using blue arrows. The medial forefoot placement is shown in the illustration.

One fresh-frozen cadaveric specimen was tested which included the first ray (distal phalanx; proximal phalanx; hallucial sesamoids; first metatarsal; medial cuneiform) and the second ray (distal phalanx; middle phalanx; proximal phalanx, second metatarsal, and; medial band of the plantar fascia). Two conditions were evaluated: (1) a planus arch-alignment and (2) a normal arch-alignment. From planus to normal posture, there was a trend toward decreased loading in both the first MTP and first MTC joints (Fig. 15.11). This finding revealed that medial forefoot joint contact mechanics may be dependent upon foot type. The cadaver model utilized a simplified set of quasi-static loads during a single phase of midstance and the flexor hallucis longus and brevis tendons were not included because static equilibrium was achieved from tension in the plantar fascia. This suggested the passive effect from the windlass mechanism was sufficient to counteract plantar loads at the hallux and metatarsals. The medial band of the plantar fascia has been shown to be critical for transmission of forces to the forefoot [27,29]. While the simulated GRF’s of the first metatarsal and hallux were vertical, a large horizontal load was present from the line-of-action of the medial band of the plantar fascia. To achieve static equilibrium in the planus model, greater tension in the plantar fascia was required to resist the combined effects of a lower first metatarsal declination angle and increased hallucial loading. These structural and functional characteristics of the planus foot appeared to promote increased first MTP and MTC joint loading. However, combining this experimental work

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FIGURE 15.11 Images of the first metatarsophalangeal joint contact pressures generated from the TekScan K-Scan 6900 pressure sensor (TekScan Inc, Boston, USA) inserted into the first metatarsophalangeal joint cavity. The pressure distributions are shown for the planus (left) and normal (right) cadaveric models.

with in silico research worked to further improve our understanding of the potential for increased stress in the first MTP and first MTC joints of the hypermobile planus foot.

15.4.4 In silico simulation Although physical testing is essential in the evaluation process, reliable computational models can augment these experiments by non-invasive predictions of biomechanical or surgical variables. It can be difficult to truly recreate the natural weightbearing environment of a joint in vitro due to soft tissue failure and constraint. FE modeling can predict soft tissue mechanics to supplement these experimental investigations. Simulations of the underlying biomechanical response of articular cartilage under weightbearing conditions may further elucidate pathways to degenerative changes. The effects of excessive first ray mobility on first MTP and first MTC joint cartilage contact mechanics are not well understood. To date, the biomechanical implications of first ray hypermobility are unknown. Therefore, we sought to understand the potential for increased stress in the first MTP and first MTC joints of the hypermobile planus foot. The cadaveric testing outlined above was utilized to calibrate, validate, and investigate the effects of first ray hypermobility on medial forefoot cartilage contact mechanics. A 3 T MRI scan (GE Healthcare, Waukesha, WI, USA) protocol was used to derive the 3D geometries of the specimen; relevant material properties were obtained from the literature. Following from calibration of the model, validation was performed for simulations of planus and non-hypermobile normal foot types during late stance in addition to simulations of first MTP, first MTC, and second MTP joint stress in the presence of first ray hypermobility. The quasi-static simulations were driven with foot-type specific physiological plantar loading and muscular forces obtained from in vivo measurements and the literature. The primary flexor and extensor tendons of the toes were added and modeled as linear elastic 1D tension-only connector elements (T3D2). The model simulations explored three loading scenarios: 70%, 80%, and 90% of the stance phase of gait. These components corresponded to the peak plantar forces in the medial forefoot and peak loading of the peroneus longus. Simulation of first ray hypermobility in the planus foot was defined by a reduction in peroneus force by 13.7% [27]. Force parameters for the plantar fascia were derived from cadaveric assessments in the literature and scaled to the weight of each subject [27] (Figs. 15.12 and 15.13). The increased passive tension theorized to occur in the plantar fascia of the planus hypermobile individual was represented by adding an initial strain of 5%. The in vivo first MTP joint angle had a mean difference in the transverse plane of 10% for the planus simulation and 7% for the normal simulation. The mean difference of the frontal plane first MTP joint angle was 4% and 6% for the planus and normal models, respectively. Compressive force in the first MTP joint was B16% higher than the first MTC joint and B41% higher than the second MTP joint. Shear forces in the superior-inferior directions were greatest in the first MTC joint by B122% compared to the first MTP joint and B117% compared to the second MTP joint. Peak superior-inferior shear force occurred at 80% of stance. There were no substantial differences in medial-lateral forces for any of the joints across foot type. The first MTC joint underwent medially directed shear forces while the

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FIGURE 15.12 A G (A) Sagittal kinematics of planus and normal feet [81] used to define the angles of the first and second metatarsophalangeal joints during stance; (B) muscle [30]; and (C) plantar fascia forces [27] used to drive the model. Plantar forces (N) derived from the planus and normal subjects in vivo, for the (D) hallux; (E) second toe; (F) first metatarsal head; and (G) second metatarsal head. (A) Adapted from Buldt AK, Levinger P, Murley GS, Menz HB, Nester CJ, Landorf KB. Foot posture is associated with kinematics of the foot during gait: a comparison of normal, planus and cavus feet. Gait Posture 2015;42(1):42 8; (B) adapted from Aubin PM, Whittaker E, Ledoux WR. A robotic cadaveric gait simulator with fuzzy logic vertical ground reaction force control. IEEE Trans Robotics 2012; (C) adapted from Erdemir A, Hamel AJ, Fauth AR, Piazza SJ, Sharkey NA, Dynamic loading of the plantar aponeurosis in walking. J Bone Jt Surg Ser A 2004.

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FIGURE 15.13 Tendon positions within the Abaqus medial forefoot FE model. The 2D connectors used to represent the tendons are highlighted in red.

first and second MTP joints demonstrated laterally oriented shear forces. The highest von Mises stress in the first MTP joint was predicted to occur during 80% of stance in the planus foot. This magnitude was B39% higher for the planus compared to normal simulation and reached 6.5 MPa. Different stress patterns were observed at the articular surface of the first metatarsal head, between foot type, but were distributed at the dorsal aspect of the joint for both simulations (Fig. 15.14). Pathologic examination of hallux rigidus has shown chondral defects to frequently occur at the dorsal apex of the dome of the first metatarsal head and adjacent to the dorsal lip of the base of the proximal phalanx [82] (Fig. 15.15). The present simulations demonstrated focal areas of stress at the dorsal aspect of first metatarsal head cartilage in the hypermobile planus simulation. Furthermore, peak first MTP joint stress occurred at the osteochondral interface of the first metatarsal head, which would suggest that degenerative changes may be initiated at the chondral surface. Several investigators have found chondral surface damage to be present in the early stages of cartilage injury mechanisms [83 90]. Increased stress at the first MTP joint in the hypermobile planus simulation may be the effect of decreased plantar load beneath the first metatarsal head and increased load beneath the hallux. A higher flexion moment arm between the hallucial load and first MTP joint likely produced focal areas of articular contact, exposing the first MTP joint to higher magnitudes of tensile and shear stress. While the current data may provide a biomechanical explanation for first ray hypermobility as a potential pathway to onset and development of hallux rigidus, future research with a longitudinal dataset is required.

15.5

Future biomechanical research

The topics covered in this chapter are broad and have each, for the most part, been extensively researched and developed. One area of future research is a comprehensive study of multiple patient populations with each of the four components described here. Relatively little is understood about how the joints of the foot work in terms of cartilage stress-strain responses to weight-bearing load and the implications these biomechanical factors may have on certain pedal pathologies. One example is the topic of hallux rigidus and first ray hypermobility, which has been explored in depth by our team with epidemiological studies [62,75,91], in vivo experimentation [64], in vitro experimentation [26], and FE modeling [63]; some work published and others currently under peer-review. Extending the integration of laboratories to other topics will serve to develop more in depth, specialized knowledge within research teams and may be considered an important method to inform future research design with a stronger evidential basis. This approach of integrating laboratories with epidemiology, in vivo, in vitro, and in silico methodologies is useful for providing

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FIGURE 15.14 Predicted von Mises stress distributions at the articular surface of the first metatarsal head. The FE images are separated by planus and normal simulations across late stance.

FIGURE 15.15 Anatomical geometries of the first metatarsophalangeal joint from MRI imaging, presenting with chondral defect at the dorsal apex of the dome of the first metatarsal head and adjacent to the dorsal lip of the base of the proximal phalanx. An image of the corresponding cadaveric cartilage is also shown.

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mechanistic insight into the onset and progression of neuromusculoskeletal disease. Such insight can spawn future studies that focus upon improved treatment strategies, tools to personalize treatment planning, and ultimately prevent disease or injury.

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Non-linear finite element model to assess the effect of tendon forces on the foot-ankle complex. Med Eng Phys 2017. [57] Anderson AE, Ellis BJ, Maas SA, Peters CL, Weiss JA. Validation of finite element predictions of cartilage contact pressure in the human hip joint. J Biomech Eng 2008. [58] Yeni YN, Christopherson GT, Dong XN, Kim DG, Fyhrie DP. Effect of microcomputed tomography voxel size on the finite element model accuracy for human cancellous bone. J Biomech Eng 2005. [59] Viceconti M, Olsen S, Nolte LP, Burton K. Extracting clinically relevant data from finite element simulations. Clin Biomech 2005. [60] Wong DWC, Wang Y, Chen TLW, Leung AKL, Zhang M. Biomechanical consequences of subtalar joint arthroereisis in treating posterior tibial tendon dysfunction: a theoretical analysis using finite element analysis. Comput Methods Biomech Biomed Eng 2017. [61] Fung A, Loundagin LL, Edwards WB. Experimental validation of finite element predicted bone strain in the human metatarsal in J Biomech 2017. [62] Morgan OJ, Hillstrom HJ, Ellis SJ, Golightly YM, Russell R, Hannan MT et al. Osteoarthritis in England: incidence trends from national health service hospital episode statistics. ACR Open Rheumatol 2019;493 8. [63] Morgan OJ, Hillstrom R, Turner R, Day J, Thaqi I, Caolo K et al. Is the planus foot type associated with first ray hypermobility?. Foot Ankle Orthop 2022.

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[64] Morgan OJ, Hillstrom R, Turner R, Day J, Thaqi I, Caolo K et al. Comparative reliability of a novel electromechanical device and handheld ruler for measuring first ray mobility. Foot Ankle Int 2021. [65] Jack EA. The aetiology of hallux rigidus in Br J Surg 1940. [66] Roukis TS, Scherer PR, Anderson CF. Position of the first ray and motion of the first metatarsophalangeal joint in J Am Podiatric Med Assoc 1996. [67] Roddy E, Menz HB. Foot osteoarthritis: latest evidence and developments. Therap Adv Musculoskelet Dis 2018. [68] Menz HB, Roddy E, Marshall M, Thomas MJ, Rathod T, Myers H, et al. Demographic and clinical factors associated with radiographic severity of first metatarsophalangeal joint osteoarthritis: cross-sectional findings from the clinical assessment study of the foot in Osteoarthr Cartil 2015. [69] Kaufman KR, Brodine SK, Shaffer RA, Johnson CW, Cullison TR. The effect of foot structure and range of motion on musculoskeletal overuse injuries in Am J Sports Med 1999. [70] Yates B, White S. The incidence and risk factors in the development of medial tibial stress syndrome among naval recruits in Am J Sports Med 2004. [71] Dixon S, Nunns M, House C, Rice H, Mostazir M, Stiles V, et al. Prospective study of biomechanical risk factors for second and third metatarsal stress fractures in military recruits in J Sci Med Sport 2019. [72] Ledoux WR, Shofer JB, Ahroni JH, Smith DG, Sangeorzan BJ Boyko EJ. Biomechanical differences among pes cavus, neutrally aligned, and pes planus feet in subjects with diabetes. Foot Ankle Int 2003. [73] Rao S, Song J, Kraszewski A, Backus S, Ellis SJ, Md JTD et al. The effect of foot structure on 1st metatarsophalangeal joint flexibility and hallucal loading. Gait Posture 2011. [74] Dieppe PA Lohmander LS, Pathogenesis and management of pain in osteoarthritis. Lancet 2005. [75] Golightly YM, Hannan MT, Nelson AE, Hillstrom HJ, Cleveland RJ, Kraus VB et al. Relationship of joint hypermobility with ankle and foot radiographic osteoarthritis and symptoms in a community-based cohort. Arthritis Care Res 2019. [76] Cowie S, Parsons S, Scammell B, McKenzie J. Hypermobility of the first ray in patients with planovalgus feet and tarsometatarsal osteoarthritis. Foot Ankle Surg 2012. [77] Doty JF, Coughlin MJ, Hirose C, Stevens F, Schutt S, Kennedy M et al. First metatarsocuneiform joint mobility: radiographic, anatomic, and clinical characteristics of the articular surface. Foot Ankle Int 2014. [78] King DM, Toolan BC. Associated deformities and hypermobility in hallux valgus: an investigation with weightbearing radiographs. Foot Ankle Int 2004. [79] Golightly YM, Hannan MT, Dufour AB, Renner JB, Jordan JM. Factors associated with hallux valgus in a community-based cross-sectional study of adults with and without osteoarthritis. Arthritis Care Res 2015. [80] Nguyen USDT, Hillstrom HJ, Li W, Dufour AB, Kiel DP, Procter-Gray E et al. Factors associated with hallux valgus in a population-based study of older women and men: the MOBILIZE Boston Study. Osteoarthr Cartil 2010. [81] Buldt AK, Levinger P, Murley GS, Menz HB, Nester CJ, Landorf KB. Foot posture is associated with kinematics of the foot during gait: a comparison of normal, planus and cavus feet. Gait Posture 2015;42(1):42 8. [82] McMaster MJ. The pathogenesis of hallux rigidus in J Bone Jt Surg Ser B 1978. [83] Brown TD, Pope DF, Hale JE, Buckwalter JA, Brand RA. Effects of osteochondral defect size on cartilage contact stress in J Orthopaedic Res 1991. [84] Flachsmann ER, Broom ND, Oloyede A. A biomechanical investigation of unconstrained shear failure of the osteochondral region under impact loading. Clin Biomech 1995. [85] Guettler JH, Demetropoulos CK, Yang KH, Jurist KA. Osteochondral defects in the human knee: Influence of defect size on cartilage rim stress and load redistribution to surrounding cartilage in Am J Sports Med 2004. [86] Hughston JC, Hergenroeder PT, Courtenay BG. Osteochondritis dissecans of the femoral condyles in J Bone Jt Surg Ser A 1984. [87] McCarthy JC, Lee JA. Acetabular dysplasia: a paradigm of arthroscopic examination of chondral injuries. Clin Orthop Relat Res 2002. [88] Messner K, Maletius W. The long-term prognosis for severe damage to weight-bearing cartilage in the knee: a 14-year clinical and radiographic follow-up in 28 young athletes. Acta Orthop Scand 1996. [89] Schenck RC, Athanasiou KA, Constantinides G, Gomez E. A biomechanical analysis of articular cartilage of the human elbow and a potential relationship to osteochondritis dissecans. Clin Orthop Relat Res 1994. [90] Tannast M, Goricki D, Beck M, Murphy SB, Siebenrock KA. Hip damage occurs at the zone of femoroacetabular impingement. Clin. Orthop Relat Res 2008. [91] Flowers P, Nelson AE, Hannan MT, Hillstrom HJ, Renner JB, Jordan JM et al. Foot osteoarthritis frequency and associated factors in a community-based cross-sectional study of White and African American adults. Arthritis Care Res. 2021.

Chapter 16

Radiographs Morgan Leslie1 and William R. Ledoux1,2,3 1

RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 2Department of

Mechanical Engineering, University of Washington, Seattle, WA, United States, 3Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Abstract Radiography, which involves the generation of 2D static images using X-ray radiation, is the oldest and simplest medical imaging technique for the diagnosis of various diseases. The technology has evolved from analog devices (some still in use) to sophisticated digital systems, which are increasingly wide spread. Even though there are many specialized radiographic views in the context of the foot and ankle, there are three primary ones: anterior/posterior, medial/lateral, and hindfoot alignment. Similarly, there are numerous measurements that can made on the bones of the feet, with some important ones that describe overall foot morphology including: in the sagittal plane: calcaneal pitch angle, navicular height, and the lateral talometatarsal angle (all measures of arch height); in the transverse plane: the anteroposterior talometatarsal angle (a measure of forefoot abduction); and in the hindfoot alignment view: the tibial-calcaneal distance, a measure of hindfoot eversion. There are also many foot-specific considerations, some of which are reviewed here. Finally, a representative group of X-ray measures of foot pathology are discussed, and some of the basic concerns with using radiographs are reviewed.

16.1

Introduction

Radiographs are essentially the generation of still two-dimensional (2D) images using X-ray energy. X-ray photons are emitted and aimed at a part of the body; some are absorbed by tissue, others are scattered, and some penetrate through and can be captured, either by film or digital devices. These captured X-rays form a shadow image of the energy that is absorbed by the body. In their simplest form, static 2D images can used to visualize dense tissues, primarily bones. More complex uses of X-ray energy for imaging include three-dimensional (3D) computed tomography (CT) and dynamic analyses via fluoroscopy. This chapter briefly discusses the use of static X-ray images (or radiographs) to study the foot and ankle. Radiographs of the foot have been used diagnostically since at least the mid-19th century. Beal and Grant used coronal plane images of the foot to explore limb length discrepancies [1]. Cahoon presented a full set of X-ray views that could be used for diagnosing foot pathology [2]. Gamble suggested how foot X-rays could be used to evaluate the foot as a “mechanism” [3]. These early papers helped lay the ground work for today’s standard radiographic analyses of the foot.

16.2

Radiographic technology

Analog radiography was prevalent in the earliest X-ray devices and is still in use today. These systems consist of a cassette containing an X-ray film that stores a latent image upon X-ray exposure. This is then processed in a darkroom to generate the radiograph [4]. Digital radiography (DR) requires no film and can display the radiograph image directly on a computer monitor [4]. In the 1980s, advances in digital technologies brought DR methods into the commercial market and they have become the standard in recent years [5]. The most commonly used DR systems in clinical practice are computed radiography (CR), direct detection flat panel systems, indirect detection flat panel systems, and Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00055-X © 2023 Elsevier Inc. All rights reserved.

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charged-coupled devices (CCDs) [4,6]. With CR, an imaging plate made of barium fluorobromide or barium fluoride traps the X-ray beam within the phosphor layer, which is then processed by a laser that releases the stored electromagnetic energy as light that can be read by a photodiode [6]. The imaging plate must be processed rapidly because the latent image can only be stored on the plate for minutes—approximately 25% of the stored energy is lost within 8 hours of exposure [4]. Indirect and direct flat panel DR systems differ in the way they convert the X-rays. Direct DR converts X-rays to an electrical signal using a photoconductive layer, while indirect DR converts X-rays first to light and then to an electrical signal using a scintillator made up of cesium iodide crystals [4]. Both systems fit into an X-ray table and are connected to a computer with a cable to view the image [4]. CCDs consist of a scintillator that converts the X-rays into light and an optical device that minimizes and focuses the light which is then stored as a pixel matrix on a CCD chip that converts the light to an electrical signal [4].

FIGURE 16.1 (A) Well-positioned anteroposterior view foot radiograph. The red circle marks the third cuneiform and shows where the central beam should be aimed. (B) The cassette is placed on the floor. The patient stands and places their foot on the center of the cassette. The knee should be fully extended. The red circle marks the third cuneiform and shows where the central beam should be aimed. The central beam should be angled 15 degrees proximally [10]. CR, computed radiography. Copyright 2016, Rubin Institute for Advanced Orthopedics, Sinai Hospital of Baltimore.

FIGURE 16.2 (A) Well-positioned lateral view foot radiograph. The red circle marks the third cuneiform and shows where the central beam should be aimed. (B) The central beam is perpendicular to the cassette. While the patient is weight-bearing, the lateral aspect of the foot is positioned against the cassette. Note the neutral positioning of the ankle and that the patient has taken a small step backward with the contralateral foot. The red circle marks the third cuneiform and shows where the central beam should be aimed [10]. CR, computed radiography. Copyright 2016, Rubin Institute for Advanced Orthopedics, Sinai Hospital of Baltimore.

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267

Standard radiographic views of the foot and ankle

Numerous foot and ankle X-ray views have been proposed in the literature, often for very specialized purposes. For instance, the ankle mortise view, which is a modified coronal plane view with the foot in 20 degrees of internal rotation [7], has been used to determine ankle alignment [8] and the oblique view, which has the patient supine on a table with the knee flexed and the lateral border of the foot raised 30 degrees, has been used to determine forefoot alignment [9]. However, this chapter mainly considers the three radiographic views that provide most of the widely used radiographic measures of foot shape: anterior/posterior (Fig. 16.1), medial/lateral (Fig. 16.2), and hindfoot alignment (Fig. 16.3). Lamm et al. have provided an excellent review of these radiographic views and the subsequent measurements [10].

FIGURE 16.3 (A) Well-positioned hindfoot alignment view radiograph. Note that both distal tibiae and feet are included. The red circle should be placed between the ankle joints and shows where the central beam should be aimed. (B) The central beam is directed at the level of the ankle joints. The patient stands on an elevated Plexiglass radiograph box. The cassette should be angled 20 degrees and placed in a slot that is in front of the patient. The central beam should be aimed in a posterior anterior orientation between the ankle joints. The beam should be perpendicular to the radiographic cassette (i.e., 20 degrees from the surface of the Plexiglass box) [10]. CR, computed radiography. Copyright 2016, Rubin Institute for Advanced Orthopedics, Sinai Hospital of Baltimore.

FIGURE 16.4 (A) Normal sagittal plane measurements and reference points. (B) Normal sagittal plane measurements. DPHA, dorsal proximal hallux angle; DPMA, dorsal proximal metatarsal angle; DPPA, dorsal proximal phalangeal angle; PDHA, plantar distal hallux angle; PDMA, plantar distal metatarsal angle; SD, standard deviation [10]. Copyright 2016, Rubin Institute for Advanced Orthopedics, Sinai Hospital of Baltimore.

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FIGURE 16.5 (A) Normal anteroposterior view measurements. IMA, intermetatarsal angle; JLCA, joint line convergence angle; LDHA, lateral distal hallux angle; LDMA, lateral distal metatarsal angle; MPHA, medial proximal hallux angle; MPMA, medial proximal metatarsal angle; MPPA, medial proximal phalangeal angle; SD, standard deviation; TCA, talocalcaneal angle. (B) Metatarsus adductus angle (MAA) shows the angular relationship of the midfoot with respect to the second metatarsal mid-diaphyseal line. (C) The normal mechanical axis deviation (MAD) is 4 mm (range 2 6 mm) lateral. (D) Anteroposterior view measurements. AP, anterior/posterior; HAA, hallux abductus angle; HIA, hallux interphalangeal angle [10]. Copyright 2016, Rubin Institute for Advanced Orthopedics, Sinai Hospital of Baltimore.

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Each of these figures demonstrates the position of the patient and X-ray system, as well as where the central beam should be located.

16.4

Definitions of X-ray measurements of foot shape

From the three emphasized radiographic views (Figs. 16.1 16.3), it is possible to make numerous measurements. The readers are recommended to see Lamm et al.’s excellent review for the details of how to make each measurement (Figs. 16.4 16.6) and for representative normal values [10]. The primary measures of foot shape or arch

FIGURE 16.6 (A) A bisector line drawn halfway between the medial and lateral aspects of the trochlea of the talus marks the normal center of the talar dome. The plafond malleolar angle (PMA) shows the relationship between the tibial plafond and the transmalleolar axis. The transmalleolar axis is drawn from the tip of the medial malleolus to the tip of the lateral malleolus. (B) Normal axial view measurements and reference points for the axial view joint line convergence angle (JLCA) and foot height [10]. Copyright 2016, Rubin Institute for Advanced Orthopedics, Sinai Hospital of Baltimore.

FIGURE 16.7 Lateral weight-bearing radiograph, marked with different geometric angles used for diagnoses: (A) Lateral talometarsal (Meary’s) angle, (B) Dijian-Annonier angle (less than 120 degrees, indicating a cavus foot), (C) calcaneal pitch angle, (D) Talocalcaneal angle, (E) Hibb’s angle [17].

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TABLE 16.1 Common foot arch abnormalities and their related clinical radiographic measurements [18]. Geometric measurement (based on radiograph view) Deformity

AP View

Pes Planus

G

Talonavicular coverage angle greater than 7 degrees

Lateral View G

G

Pes Cavus

G

G

Meary’s angle greater than 4 degrees convex downward (greater than 30 degrees is considered severe) Calcaneal pitch less than “normal range” of 17 to 32 degrees. Meary’s angle greater than 4 degrees convex upward (greater than 30 degrees is considered severe) Calcaneal pitch angle greater than “normal range” of 17 to 32 degrees.

height in the sagittal plane (Fig. 16.4) include the calcaneal pitch angle (or calcaneal inclination angle), the navicular height, and the lateral talometatarsal angle (or lateral Meary’s angle). In the transverse plane (Fig. 16.5), an important measure of foot shape or forefoot abduction is the anteroposterior talometatarsal angle (or AP Meary’s angle). Finally, in the hindfoot alignment view (Fig. 16.6), the tibial-calcaneal distance represents the amount of hindfoot eversion.

16.5

Foot-specific applications and considerations

The foot contains intricate 3D interlocking bony structures, and it can be difficult to capture adequate information in a 2D radiographic image due to overlapping bony features in many views. However, radiography is considered as the best method for diagnosing foot and ankle disorders in terms of cost, speed, and ease of use [11,12]. Many radiographic textbooks describe foot examination techniques, but there appears to be a great degree of variation from these principles in practice [11]. A review of the various techniques is beyond the scope of this chapter, but a few representative considerations are presented. One study used anthropometric foot phantoms to assess image quality from employing different cranial central beam angulations showed that, in agreeance with textbook guidelines, 15 degrees of cranial angulation produced the overall best quality X-ray images and 0 degrees of angulation was optimal for oblique images of the entire foot [11]. The amount of weight-bearing should also be considered [11], as placing the foot under load will not only alter the foot arches, but also potentially increase the tarsometatarsal joint spaces [13,14] and the hallux valgus angle (HVA) [15]. Willauer et al. examined the effect of misalignment on both anterior/posterior and medial/lateral foot radiographs [16]. They found that certain parameters, for example, calcaneal pitch, were more sensitive to X-ray misalignment, while others were more robust and less sensitive. A study compared two different hindfoot alignment methods and determined that “Buck’s” method is able to obtain quality images more quickly and easily than “Cobey’s” method without any negative implications on the clinical evaluation [12].

16.6

Clinical X-ray measures of foot shape

Foot and ankle radiographs are essential in identifying deformities and to help inform decisions as to whether a correction surgery might be required. Geometric measurements can be taken from the radiographs to identify different deformities and quantify their severities. Two different angles in the lateral weight-bearing view have been widely used for the diagnosis of a pes cavus deformity: the Dijan-Annonier angle (at the medial arch) and the Hibb’s angle (between the long axis of the calcaneus and the first metatarsal) [17]. A Dijian-Annonier angle less than 120 degrees and a Hibb’s angle of more than 45 degrees indicate the presence of a cavus foot (Fig. 16.7) [17]. These angles can also provide insight for the diagnoses of different foot arch abnormalities, such as flatfoot (pes planus) or high arch (pes cavus) (Table 16.1).

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Another radiograph-derived measurement, the metatarsus primus elevatus (MPE), has been debated as an indicative factor of hallux rigidus (HR), and it is considered as a way to provide a threshold quantitative value to diagnose this deformity in a lateral weight-bearing radiograph [19]. There has been debate on this correlation, mostly due to the lack of a proven threshold value to define a “gold-standard” pathologic value in addition to lack of precision and/or accuracy in radiographs [19]. Bouaicha et al. developed a reliable and validated method to measure MPE to determine a threshold value (greater than 5.0 mm) for identification of HR (Fig. 16.8) [19]. Lateral weight-bearing radiographs are not the only clinically useful view of the foot and ankle for providing diagnostic measurements. For measurements of the HVA and intermetatarsal angle (IMA), weight-bearing anterior-posterior radiographic views of the foot are taken (Fig. 16.9). Non-weight-bearing radiographs tend to overestimate these measurements in early incidences of hallux valgus (HV) and under-estimate them in advanced progressions of HV [20]. An average HVA angle greater than 15 degrees (with greater than 40 degrees denoting severe cases) and an IMA greater than 9 degrees (with greater than 18 degrees denoting severe cases) indicates the presence of HV deformity [21]. For diagnoses specific to the ankle or subtalar joint, the mortise view and lateral view weight-bearing radiographs are mostly used [22]. The mortise view consists of 10 degrees internal rotation of the foot and ankle, with even weight

FIGURE 16.8 Radiographic assessment of metatarsus primus elevatus (MPE). A circle (C) is fit onto the metatarsal head and a tangent (T) is drawn along the dorsal cortical bone of the first metatarsal shaft. Where the circle intersects the tangent line most proximally, a perpendicular line is drawn through the dorsal cortex of the second metatarsal shaft. The distance (X) is measured along that perpendicular line between the two intersection points of the two dorsal cortical bones. This measurement is the MPE. (A) represents an illustration of this method, where (B) shows a clinical example of the same method on a lateral weight-bearing radiograph [19].

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FIGURE 16.9 The measurement method of the hallux valgus angle (HVA) and the intermetarsal angle (IMA) in an anterior-posterior weight-bearing radiograph of a foot with hallux valgus (HV) deformity [20].

distribution on both legs [22]. (Listed as 20 degrees internal rotation elsewhere [7].) Lateral weight-bearing radiographs center the X-ray beam on the medial aspect of the ankle, with the foot and ankle in 15 degrees of external rotation [22]. Osteoarthritis of the tibiotalar joint and the subtalar joint can be identified and graded (Kellgren-Lawrence) by analysis of mortise view and lateral view radiographs for presence and severity of joint space narrowing and osteophytes (Figs. 16.10 and 16.11) [22].

16.7

Issues with X-ray measures of foot shape

While ease of use and the relatively inexpensive nature of radiographic hardware (compared to CT or magnetic resonance imaging (MRI) scanners) make radiographs a popular imaging modality, numerous issues should be noted. First, every X-ray image exposes the patient to radiation and increases their long-term risk of cancer. This is mitigated somewhat by the use of collimation and shielding. Second, errors in radiographic technique (e.g., misaligning the central beam, incomplete weight-bearing, or incorrect cassette placement) can led to incorrect images, and therefore, incorrect measurements. Careful training of X-ray technicians is required to prevent this. Third, radiographs are always 2D projections of 3D objects, and they are always magnified to some degree depending how far the object to be imaged is relative to the cartridge or detector. This requires an understanding of the 2D/3D relationship and careful placement of the foot. Generally, all of these issues can be sufficiently addressed with forethought and adequate training.

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FIGURE 16.10 Mortise view of the ankle joint: (A) normal, grade 0, (B) osteophyte at medial tibia, grade 1, (C) osteophyte at medial tibia, grade 2, (D) osteophyte at medial tibia, grade 3 [22].

16.8

Areas of future biomechanical research

While X-ray imaging is a mature technology that is well established and validated in the context of the foot and ankle, there are still areas of potential innovation. Obtaining many of the measurements described earlier in this chapter tends to be a manual task that requires the operator to identify anatomical landmarks, resulting in the potential errors and biases associated with this. Machine learning and related techniques have been explored for medical image analysis and could potentially be applied to automate the analysis of the foot and ankle measurements, increasing reliability and reducing the time needed to obtain the information.

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FIGURE 16.11 Lateral view of the subtalar joint (STJ): (A) normal, grade 0, (B) joint space narrowing, grade 1, (C) joint space narrowing, grade 2, (D) joint space narrowing, grade 3 [22].

A new imaging method called the biplanar linear EOS system has recently been developed and investigated to determine its advantages and utility as an alternative to conventional radiographs [5]. There have been studies confirming that measurements in the biplanar system are similar or superior to those from conventional radiographs for sagittal images of the spine and pelvis, total hip arthroplasty pre-operative planning, and identifying lower limb length discrepancies [5].

References [1] Beal MC, Grant JH. Standing foot x-rays. J Am Osteopath Assoc 1947;46(5):306. [2] Cahoon Jr. JB. Radiography of the foot. Radiography 1949;15(174):129 36. [3] Gamble FO. Foot mechanics and the x-ray. J Am Podiatr Assoc 1959;49(4):168 71.

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[4] Drost WT. Transitioning to digital radiography. J Vet Emerg Crit Care (San Antonio) 2011;21(2):137 43. [5] Chua C, Tan S, Lim A, Hui J. Accuracy of biplanar linear radiography vs conventional radiographs when used for lower limb and implant measurements. Arch Orthop Trauma Surg 2021;142:735 45. [6] Hon D, Carrino J. Radiography. In: Waldman S, Bloch J, editors. Pain management. W.B. Saunders; 2007. p. 74. [7] Barg A, Harris MD, Henninger HB, Amendola RL, Saltzman CL, Hintermann B, et al. Medial distal tibial angle: comparison between weightbearing mortise view and hindfoot alignment view. Foot Ankle Int 2012;33(8):655 61. [8] Wang B, Saltzman CL, Chalayon O, Barg A. Does the subtalar joint compensate for ankle malalignment in end-stage ankle arthritis? Clin Orthop Relat Res 2015;473(1):318 25. [9] Lee KT, Kim KC, Park YU, Kim TW, Lee YK. Radiographic evaluation of foot structure following fifth metatarsal stress fracture. Foot Ankle Int 2011;32(8):796 801. [10] Lamm BM, Stasko PA, Gesheff MG, Bhave A. Normal foot and ankle radiographic angles, measurements, and reference points. J Foot Ankle Surg 2016;55(5):991 8. [11] Flintham K, Snaith B, Field L. Review and optimisation of foot radiography technique. Radiography (Lond) 2021;27(2):284 8. [12] Yang C, Xu X, Hu M, Wang B, Zhu Y, Liu J. Optimization of hindfoot alignment radiography. Acta Radiol 2017;58(6):719 25. [13] Shereff MJ, DiGiovanni L, Bejjani FJ, Hersh A, Kummer FJ. A comparison of nonweight-bearing and weight-bearing radiographs of the foot. Foot Ankle 1990;10(6):306 11. [14] Fuhrmann RA, Layher F, Wetzel WD. Radiographic changes in forefoot geometry with weightbearing. Foot Ankle Int 2003;24(4):326 31. [15] Tanaka Y, Takakura Y, Takaoka T, Akiyama K, Fujii T, Tamai S. Radiographic analysis of hallux valgus in women on weightbearing and nonweightbearing. Clin Orthop Relat Res 1997;(336):186 94. [16] Willauer P, Sangeorzan BJ, Whittaker EC, Shofer JB, Ledoux WR. The sensitivity of standard radiographic foot measures to misalignment. Foot Ankle Int 2014;35(12):1334 40. [17] Maynou C, Szymanski C, Thiounn A. The adult cavus foot. EFORT Open Rev 2017;2(5):221 9. [18] Escobedo E., Pinney S., Hunter J., Sangeorzan B. Evaluation of adult foot alignment; 2016 ,http://uwmsk.org/footalignment/doku.php.. [19] Bouaicha S, Ehrmann C, Moor BK, Maquieira GJ, Espinosa N. Radiographic analysis of metatarsus primus elevatus and hallux rigidus. Foot Ankle Int 2010;31(9):807 14. [20] Boszczyk A, Kwapisz S, Kicinski M, Kordasiewicz B, Liszka H. Non-weightbearing compared with weightbearing x-rays in hallux valgus decision-making. Skelet Radiol 2020;49(9):1441 7. [21] Pique-Vidal C, Vila J. A geometric analysis of hallux valgus: correlation with clinical assessment of severity. J Foot Ankle Res 2009;2:15. [22] Kraus VB, Kilfoil TM, Hash 2nd TW, McDaniel G, Renner JB, Carrino JA, et al. Atlas of radiographic features of osteoarthritis of the ankle and hindfoot. Osteoarthr Cartil 2015;23(12):2059 85.

Chapter 17

Computed Tomography of the Foot and Ankle Scott Telfer1,2,3, Christina L. Brunnquell4 and William R. Ledoux1,2,3 1

Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States, 2Department of Mechanical Engineering,

University of Washington, Seattle, WA, United States, 3RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 4Department of Radiology, University of Washington, Seattle, WA, United States

Abstract Computed tomography (CT) is a medical imaging modality that has found a range of clinical and research applications related to foot and ankle biomechanics. Data from a CT scan can provide detailed information about the anatomy of the foot and ankle in health and disease, particularly in relation to bone. This makes CT an essential technology for tasks such as fracture assessment, study of bone morphology, and computation model development. In this chapter, we give an overview of the technology and its development, review its applications in the context of foot and ankle biomechanics, and propose future areas of research.

17.1

Introduction

Computed tomography (CT) plays an important role in clinical and research biomechanics. This is largely due to its ability to produce detailed cross-sectional information on the internal anatomy of a body in a noninvasive manner. In general, modern CT scanners work by rotating an X-ray source and opposing detector around the body [1]. The signal is mathematically processed to generate a cross-sectional image, or “slice,” of the body. If successive slices through the body are collected, usually by using a multirow detector or moving the patient with respect to the CT gantry, they can be “stacked” together to create a 3D volume of images that allows for improved location of the relative position of anatomical structures, identification of abnormalities, and the recreation of 3D models of the anatomy (Fig. 17.1). The detailed information that CT can provide has made it an essential tool in both the clinical and research realms. This chapter will (1) provide a brief overview of CT imaging technology, its development, and variations; (2) review how CT imaging has been applied in the context of foot and ankle biomechanics (both clinical and research); and (3) propose areas of future research where CT imaging may be further utilized.

FIGURE 17.1 (A) Stack of cross-sectional images (slices); (B) bony anatomy identified (segmented) on each image using a combination of automatic and manual techniques; (C) 3D model of anatomy of interest reconstructed from segmented data. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00011-1 © 2023 Elsevier Inc. All rights reserved.

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17.1.1 History and development of computed tomography The first-reported CT scan was carried out in 1971 by Godfrey Hounsfield [2] who later won the Nobel Prize for Physiology or Medicine for his work in developing the technology. (As an interesting side note, there is some debate as to whether or not the Beatles played a role in the development of the CT scanner [3]). Since its conception, the technology has gone through several stages of development, generally referred to as “generations,” to reach the CT scanners that are seen in hospitals across the world today [4]. Initial work using CT scanning focused on brain imaging after stroke [5], but the wide ranging utility of the technology was quickly realized. Studies reporting CT scans of the foot and ankle began to appear in the literature in the early-to-mid 1980s. These articles included initial explorations of the use of the technology to show cross-sectional anatomy, usually with corresponding anatomy from cadaver specimens also presented [6,7], fractures of the tibia [8,9] and calcaneus [10], and tarsal coalition [11]. At the time of writing, a search of Pubmed database for “foot ankle computed tomography” shows over 100 articles being published in this area each year since 2015. Variations on standard CT include dual energy CT, where the anatomy of interest is imaged with X-rays at two energy levels [12]. By comparing the levels of attenuation at these different X-ray energy ranges, distinct tissue characteristics can be identified, allowing for processed images with improved soft tissue contrast or material identification and separation [13]. Dual energy CT can also be used to determine bone mineral density [14], Achilles tendinopathy [13], and lower extremity vascular issues [15], but by far the most common use in foot and ankle imaging is to assess the presence of uric acid crystals associated with gout [16]. A further variation on CT that is becoming more widely used for foot and ankle imaging is dedicated extremity cone beam CT (CBCT, Fig. 17.2). Primarily utilized for dental imaging but gaining popularity for the foot and ankle, this technique utilizes a cone-shaped X-ray beam that covers a relatively large volume in a single rotation around the subject [17]. It has the advantages of a lower radiation dose than a standard clinical CT scan, and that the patient or subject can stand upright in the scanner and thus weight-bearing images can be obtained without complicated and uncomfortable loading apparatuses [17]. This may be important for assessing the functional alignment of the foot bones. This variation on CT imaging is discussed in greater detail elsewhere in this book.

17.1.2 Comparison to other imaging modalities Arguably, the imaging modality that is the closest alternative to CT is magnetic resonance imaging (MRI). MRI uses a magnetic field and radio waves to produce similar cross-sectional images of the anatomy. This is discussed in more detail elsewhere in this book, but in general, MRI can provide significantly improved soft tissue contrast, whereas CT provides more information on bony structures with improved spatial resolution. CT has several advantages over MRI: (1) data acquisition is considerably faster. Scans of the foot and ankle can be obtained quickly, with the present generation of scanners able to capture 1200 slices each consisting of a 512 3 512 image matrix in around 1 second [1]; (2) CT data can be captured at high spatial resolution with isotropic or near isotropic voxels for larger volumes; and (3) there are few contra-indications for CT. One of the disadvantages of CT relative to MRI is that it utilizes ionizing radiation. A “typical” clinical CT scan of the foot and ankle will result in an effective radiation dose of less than B0.1 mSv [18]. To place this figure in context, the average annual background radiation dose for an adult living in the USA is

FIGURE 17.2 (A) Standard clinical CT scanner. For a foot and/or ankle scan, the individual would lie on the bed and the bed is moved so that the anatomy of interest is passed through the gantry; (B) Cone beam CT scanner where individual stands or sits with feet and ankles in gantry. Part (A) of this figure is licensed under the Creative Commons Attribution-Share Alike 4.0 International license.

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around 3 mSv, from a combination of terrestrial, cosmic, or internal sources [19]. In addition, if contrast agents are used to help image soft tissue structures within the anatomy [20], there is some risk of an allergic reaction. Other medical imaging options that should be noted include planar X-rays and ultrasound. Planar X-rays, or radiographs, are often used for the initial assessment of foot and ankle problems and have the advantage of a lower radiation dose [21], but these projection images lack the 3D detail that CT can provide. Ultrasound imaging is also used regularly for foot and ankle imaging and can be particularly useful for soft tissue injuries, but compared to CT it has limited ability to image bone. Although there have been developments in 3D imaging using ultrasound, the capabilities are still far behind those of CT [22].

17.1.3 Computed tomography protocols for the foot and ankle For high spatial resolution and image quality in CT imaging of the foot and ankle, several patient positioning factors, acquisition parameters, and reconstruction settings can be optimized [23]. Relevant to patient positioning, spatial resolution is maximized near the center of the imaging field of view (FOV) and degradation is observed near the edges [24]. Therefore alignment of the anatomy of interest in the center of the CT FOV can improve spatial resolution. In addition, if appropriate for the imaging task, orienting the foot oblique to the axial scan plan can help reduce large angular variabilities in X-ray attenuation, thereby reducing image noise. If the imaging task or patient requires orientation of the foot within the transverse plane (orthogonal to the shin), angular tube current modulation (TCM) can be used. Finally, if the imaging task is focused on just one foot or ankle, the contralateral foot should be excluded if hardware is present. A typical foot or ankle acquisition protocol uses an X-ray tube voltage (kV) of 100 120. The detector configuration should be selected such that detector data is not binned and thin slices (on the order of 0.5 0.6 mm) are accessible for reconstruction and sagittal or coronal reformats. The minimum scan FOV should be selected. The tube current-time product (mAs) can be selected manually, or automatic TCM can be utilized; the appropriate mAs value will depend on the required image quality and the available reconstruction options. If maximizing spatial resolution is a priority for the imaging task, the user can ensure that the small X-ray focal spot (SFS) is used by selecting a manual mA less than the SFS limit for the acquisition kV. This SFS tube current limit varies by vendor and CT scanner model and should be available to the user in the scanner technical documentation. A pitch of less than 1 in helical acquisitions may help improve resolution in the slice thickness direction for tasks requiring high spatial resolution. If automatic TCM is used by default, the operator should be aware of the potential for impact of metal in the imaging field on TCM performance; manual mAs selection may be more appropriate in these situations. Reconstruction of foot and ankle acquisitions should utilize the minimum reconstruction diameter encompassing the anatomy to maximize spatial resolution in the final reconstructed image (Fig. 17.3). Various reconstruction filters (also called kernels or algorithms) exist for different purposes, such as enhancing edges, maintaining high spatial resolution, or reducing image noise. Dedicated bone algorithms or other high-resolution filters can be used to better define the boundaries of bone (Fig. 17.4). Clinically, these are particularly important for imaging fractures [25]. Metal artifact reduction (MAR) algorithms can be applied to reduce the impact of metal artifacts on the final image appearance (Fig. 17.5). However, MAR can affect surrounding CT numbers, impact the visualization of contrast enhancement, and

FIGURE 17.3 Impact of reconstruction diameter on spatial resolution. (A) A fitted 18-cm diameter reconstruction field of view and (B) a larger 30cm diameter FOV results in loss of spatial resolution. From Weigelt L, Fu¨rnstahl P, Hirsiger S, Vlachopoulos L, Espinosa N, Wirth SH. Threedimensional correction of complex ankle deformities with computer-assisted planning and patient-specific surgical guides. J. Foot Ankle Surg. 2017;56:1158 1164.

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FIGURE 17.4 Impact of CT reconstruction filter on appearance of bony structures. (A) High-resolution bone kernel and (B) soft tissue kernel applied to an axial ankle CT image.

FIGURE 17.5 Metal artifact reduction reduced the appearance of metal artifacts on reconstructed CT images. The final image appearance can vary depending on MAR algorithm (A: single energy metal artifact reduction vs B: model-based iterative reconstruction), image dose level (B: 28mGy.cm vs C: 13.6 mGy.cm vs D: 4.5 mGy.cm), and other reconstruction settings. From Grandmougin A, Bakour O, Villani N, Baumann C, Rousseau H, Teixeira PAG, Blum A. Metal artifact reduction for small metal implants on CT: which image reconstruction algorithm performs better? Eur J Radiol. 2020;127:108970.

in some circumstances introduce image artifacts, so reconstructing images both with and without the MAR algorithm on can be helpful for understanding the source of potential artifacts [23]. MAR algorithms may also have varying effects on images reconstructed with standard or soft tissue versus high-resolution reconstruction filters and the effectiveness of the algorithm can be affected by the image dose level.

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Foot-specific applications and considerations

17.2.1 Disease diagnosis CT scans of the foot and ankle are ordered by physicians for a range of reasons, including but not limited to assessment of bone fractures (covered in more detail in the following sections), identification of bone tumors [26], assessment of deformities (including developmental [27,28]), identification or localization of foreign bodies, and assessment of infections. Arthritic conditions including osteoarthritis, rheumatoid arthritis, and gout [29] are also often imaged using CT. Similarly, anatomical alterations in diabetic neuropathic feet (e.g., presence of claw toes, decreased intrinsic muscle volume, or thinner plantar aponeuroses) can be imaged with a CT scanner (Fig. 17.6) [30].

17.2.2 Surgical assessment and planning In injuries and diseases of the foot and ankle that may require surgical fixation, a CT is commonly obtained to provide the surgeon with information they can use to assess and classify the problem and plan the best repair strategy if necessary. Compared to radiographs, providing surgeons with a CT scan was found to alter the treatment of 38% of patients who had ankle fractures [31]. Plain X-rays have also been found to underestimate the size and displacement of malleolar fractures compared to CT [32], and the decision as to whether surgery is required largely driven by these measurements [33]. CT plays an important role in identifying and characterizing calcaneal fractures [34]. Postoperatively, CT is useful for monitoring many procedures [35], and has been demonstrated to be better than planar X-rays for evaluating the progression of hindfoot arthrodesis [36]. It can still be challenging to interpret fracture patterns on the individual slices of a CT scan. However, several studies have reported a statistically significant impact of the use of 3D renderings on the reproducibility of fracture classification system. A 3D model of the calcaneus improves the evaluation of calcaneal fractures, particularly for complex cases and for less experienced surgeons [37,38]. CT images have also been used for a technique for quantifying the energy associated with an ankle fracture (discuss in detail elsewhere in this book). Fracture energy correlates with

FIGURE 17.6 Technique for quantifying intrinsic muscle volume. (A) Filtered 2D slice, (B) segmented 2D slice, and (C) 3D intrinsic muscle volume.

FIGURE 17.7 Intraarticular corrective osteotomy of the medial malleolus (MM) of patient 1 with patient-specific guides. (A) Application of the osteotomy guide with the incorporated drill sleeves. (B) Completion of osteotomy with cannulated chisel. (C) Mobilization of the medial malleolus (MM). (D) Application of the reduction guide. (E) Intraoperative fluoroscopic control of temporary k-wire fixation. (F) Final screw fixation.

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FIGURE 17.8 (A) Complete FE model, (B) structure inside an encapsulated soft tissue showing bone, cartilage, fat, plantar fascia, ligament and tendon, (C) extensor apparatus of the first ray pre-surgery, (D) first ray pre-surgery, (E) first ray post-surgery modified Jones procedure, (F) first ray post-surgery FHL tendon transfer. For (C, D, E and F), tendon and extensor apparatus (blue), ligament (green), digital slips of the plantar fascia (red) and proximal (pink) and distal (yellow) cartilage are shown. Dash line indicates element passing inside or behind the bone surface. From Isvilanonda V, Dengler E, Iaquinto JM, Sangeorzan BJ, Ledoux WR. Finite element analysis of the foot: model validation and comparison between two common treatments of the clawed hallux deformity. Clin Biomech (Bristol, Avon). 2012;27:837 44. https://doi.org/10.1016/j.clinbiomech.2012.05.005.

subjective clinical ranking of severity and strongly predicts Kellgren-Lawrence scores (i.e., a measure of OA severity) [39]. CT scans acquired prior to total ankle arthroplasty can be used to design patient-specific plans and produce 3Dprinted custom guides to allow the full procedure to be optimized for the patient (Fig. 17.7, [40]). There is early evidence that this approach can provide accurate and reproducible alignments [41]. 3D-printed anatomical models can also be used to plan surgical approaches [42] and prepare surgical hardware [43]. The use of a 3D-printed model for preoperative planning of tibial fractures was found to reduce surgical times by 48% [44].

17.2.3 Biomechanics research 17.2.3.1 Kinematic measurements Studies have used CT scans of the foot and ankle that included skin markers used to track the motion of foot segments. These scans can define the bone pose and position of skin markers relative to the bone with a high degree of accuracy, helping to improve the fidelity of kinematic gait models [18,45]. Bone models produced from CT scans are also used as an input to single [46] or biplane fluoroscopy [47] based assessments of foot kinematics to determine the bone pose from 2D images. This is discussed in more detail elsewhere in the book.

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FIGURE 17.9 (Top row) Posterior view coronal plane for cavus (left), neutral aligned (middle), and planus (right) foot types. (Middle row) Superior view transverse plane for cavus (left), neutral aligned (middle), and planus (right) foot types. (Bottom row) Medial view sagittal plane for cavus (left), neutral aligned (middle), and planus (right) foot types.

17.2.3.2 Bone density properties Bone mineral density is commonly measured using dual energy absorptiometry (DEXA) and quantitative ultrasonography. These assessments are, however, generally limited to the calcaneus and provide only a gross measure of bone mineral density. In cases involving diabetic foot disease, for example, much of the damage occurs at the midfoot [48]. More detailed bone density measurements can be estimated from CT scans, estimated from the Hounsfield number of each voxel. Commean et al. used volumetric CT measurements to assess bone mineral density in people with diabetic peripheral neuropathy and healthy controls, finding that the measurements were highly reproducible between scans and raters for the metatarsal and tarsal bones [49]. Their findings demonstrated large variations within and between bones and subjects [50]. Single-slice measurements of metatarsal bone density using QCT are reliable [51].

17.2.3.3 Computational models CT scans have been extensively used to recreate bone geometry necessary for finite element models of the foot (Fig. 17.8), and are often a key part of these workflows [52 56]. This may be in combination with MRI or other modalities to obtain information about soft tissues (ligament or tendon insertion points, for example). Beyond geometry, using CT measurements of bone density to estimate the regional mechanical properties of the bone can help to improve the accuracy of these models. Typical finite element models of the foot and ankle assign a homogenous value to bone, which in may be a valid approximation depending on the questions being asked of the model [57]. However for studies requiring a more accurate study of bone behavior, assigning elementwise mechanical properties throughout the bone may be required [58].

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FIGURE 17.10 Navicular orientation relative to the talus at the talonavicular joint (TNJ), representations of each foot type are displayed in the transverse plane (superior view) and coronal plane (anterior view), red 5 talus, green 5 talar TNJ surface, gold 5 navicular TNJ surface, blue 5 navicular. (A) Pes cavus, (B) neutrally aligned, (C) asymptomatic pes planus, (D) symptomatic pes planus. Severe feet were selected for cavus and planus example. From Louie P.K., Sangeorzan B.J., Fassbind M.J., Ledoux W.R. Talonavicular joint coverage and bone morphology between different foot types. J Orthop Res 2014;32:958 66. https://doi.org/10.1002/jor.22612.

17.2.3.4 Shape modeling and assessment As discussed, CT can provide accurate 3D models of the foot and ankle bones. Standard segmentation techniques reproduce 3D bones with an accuracy of .0.6 mm [59]. Partial weight-bearing CT scans have been used to quantify foot shape (Fig. 17.9) [60,61]. Additionally, from CT-based models of individual bones, simple, linear measurements can be assessed [62 64]. More complex methodologies for analyzing shape, such as defining slice planes to consider sections of a bone or fitting regular surfaces to bone or joint surface [63 65], have also been developed. Segmented joint surfaces have also been used to quantify bone-to-bone orientation ([63], Fig. 17.10) and subluxation in the hindfoot for different foot types [63,66]. Statistical shape analysis techniques have also been developed. Incorporating methods such as principal component analysis, this type of approach can provide information about the variation in morphology between groups, that is, identifying where differences in shape occur. These types of studies have found shape differences in subtalar bones between individuals with chronic ankle instability and healthy controls [67]. Foot type (planus, neutrally aligned, and cavus) is associated with differences in the shape of the metatarsals [68] and the bones of the hindfoot (Fig. 17.11) [69]. Individuals with osteochondral defects have altered talus and medial malleolus shapes compared to healthy controls [70].

17.3

Areas of future biomechanical research

Radiation exposure remains one of the primary concerns with CT scans, and dose optimization—using the appropriate dose that allows the clinical/diagnostic task to be accomplished without excess dose—is an important goal. Researchers have studied this using a cadaver limb and found that bone pose and attached markers could be adequately identified with a scan taken using an exposure of 50mAs, one third of that of a standard clinical CT scan [18]. Recent

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Calcaneus PC1 Cavus NA APP SPP -0.06

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FIGURE 17.11 Principal component 1 (PC1) of the calcaneus with each foot type plotted. Graphical representations of minimum in red and maximum in cyan with overlap in the middle for PC1 with medial, posterior, and superior views of a left foot from top to bottom. From Moore ES, Kindig MW, McKearney DA, Telfer S, Sangeorzan BJ, Ledoux WR. Hind- and midfoot bone morphology varies with foot type and sex. J Orthop Res. 2019;37:744 759.

developments in CBCT may help to reduce the required radiation dose; however, the accuracy of the bone models produced has not been rigorously compared against standard clinical CT scans. Micro CT involves specialized scanners that can obtain images with a resolution of ,100 µm. These can provide highly detailed images of the microstructure of bone to the level of individual trabeculae [71]. This scanning technology may have interesting applications in modeling and has been demonstrated in animal models of the foot and ankle [72]. DEXA scans of foot bones might be informative for quantifying osseous changes due to diabetic foot disease. Finally, four-dimensional CT (4DCT) allows dynamic movements to be captured. The current temporal resolution of these scans is limited, but it has been demonstrated in ankle syndesmosis injuries [73].

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[31] Magid D, Michelson JD, Ney DR, Fishman EK. Adult ankle fractures: comparison of plain films and interactive two- and three-dimensional CT scans. Am J Roentgenol 1990;154:1017 23. Available from: https://doi.org/10.2214/ajr.154.5.2108536. [32] Ferries JS, DeCoster TA, Firoozbakhsh KK, Garcia JF, Miller RA. Plain radiographic interpretation in trimalleolar ankle fractures poorly assesses posterior fragment size. J Orthop Trauma 1994;8:328 31. Available from: https://doi.org/10.1097/00005131-199408000-00009. [33] Haraguchi N, Haruyama H, Toga H, Kato F. Pathoanatomy of posterior malleolar fractures of the ankle. J Bone Jt Surg 2006;88:1085 92. Available from: https://doi.org/10.2106/JBJS.E.00856. [34] Daftary A, Haims AH, Baumgaertner MR. Fractures of the calcaneus: a review with emphasis on CT. Radiographics 2005;25:1215 26. Available from: https://doi.org/10.1148/rg.255045713. [35] Limarzi GM, Scherer KF, Richardson ML, Warden DR, Wasyliw CW, Porrino JA, et al. CT and MR imaging of the postoperative ankle and foot. Radiographics 2016;36:1828 48. Available from: https://doi.org/10.1148/rg.2016160016. [36] Coughlin MJ, Grimes JS, Traughber PD, Jones CP. Comparison of radiographs and CT scans in the prospective evaluation of the fusion of hindfoot arthrodesis. Foot Ankle Int 2006;27:780 7. Available from: https://doi.org/10.1177/107110070602701004. [37] Roll C, Schirmbeck J, Mu¨ller F, Neumann C, Kinner B. Value of 3D reconstructions of CT scans for calcaneal fracture assessment. Foot Ankle Int 2016;37:1211 17. Available from: https://doi.org/10.1177/1071100716660824. [38] Brunner A, Heeren N, Albrecht F, Hahn M, Ulmar B, Babst R. Effect of three-dimensional computed tomography reconstructions on reliability. Foot Ankle Int 2012;33:727 33. Available from: https://doi.org/10.3113/FAI.2012.0727. [39] Anderson DD, Mosqueda T, Thomas T, Hermanson EL, Brown TD, Marsh JL. Quantifying tibial plafond fracture severity: absorbed energy and fragment displacement agree with clinical rank ordering. J Orthop Res 2008;26:1046 52. Available from: https://doi.org/10.1002/ jor.20550. [40] Weigelt L, Fu¨rnstahl P, Hirsiger S, Vlachopoulos L, Espinosa N, Wirth SH. Three-dimensional correction of complex ankle deformities with computer-assisted planning and patient-specific surgical guides. J Foot Ankle Surg 2017;56:1158 64. Available from: https://doi.org/10.1053/j. jfas.2017.05.025. [41] Hsu AR, Davis WH, Cohen BE, Jones CP, Ellington JK, Anderson RB. Radiographic outcomes of preoperative CT scan derived patientspecific total ankle arthroplasty. Foot Ankle Int 2015;36:1163 9. Available from: https://doi.org/10.1177/1071100715585561. [42] Giovinco NA, Dunn SP, Dowling L, Smith C, Trowell L, Ruch JA, et al. A novel combination of printed 3-dimensional anatomic templates and computer-assisted surgical simulation for virtual preoperative planning in charcot foot reconstruction. J Foot Ankle Surg 2012;51:387 93. Available from: https://doi.org/10.1053/j.jfas.2012.01.014. [43] Chung KJ, Hong DY, Kim YT, Yang I, Park YW, Kim HN. Preshaping plates for minimally invasive fixation of calcaneal fractures using a real-size 3D-printed model as a preoperative and intraoperative tool. Foot Ankle Int 2014;35:1231 6. Available from: https://doi.org/10.1177/ 1071100714544522. [44] Corona PS, Vicente M, Tetsworth K, Glatt V. Preliminary results using patient-specific 3d printed models to improve preoperative planning for correction of post-traumatic tibial deformities with circular frames. Injury 2018;49:S51 9. Available from: https://doi.org/10.1016/j. injury.2018.07.017. [45] Oosterwaal M, Telfer S, Torholm S, Carbes S, van Rhijn LW, Macduff R, et al. Generation of subject-specific, dynamic, multisegment ankle and foot models to improve orthotic design: a feasibility study. BMC Musculoskelet Disord 2011;12:256. Available from: https://doi.org/ 10.1186/1471-2474-12-256. [46] Fukano M, Kuroyanagi Y, Fukubayashi T, Banks S. Three-dimensional kinematics of the talocrural and subtalar joints during drop landing. J Appl Biomech 2014;30:160 5. Available from: https://doi.org/10.1123/jab.2012-0192. [47] Iaquinto JM, Kindig MW, Haynor DR, Vu Q, Pepin N, Tsai R, et al. Model-based tracking of the bones of the foot: a biplane fluoroscopy validation study. Comput Biol Med 2018;92:118 27. Available from: https://doi.org/10.1016/j.compbiomed.2017.11.006. [48] Schon LC, Weinfeld SB, Horton GA, Resch S. Radiographic and clinical classification of acquired midtarsus deformities. Foot Ankle Int 1998;19:394 404. Available from: https://doi.org/10.1177/107110079801900610. [49] Commean PK, Kennedy JA, Bahow KA, Hildebolt CF, Liu L, Smith KE, et al. Volumetric quantitative computed tomography measurement precision for volumes and densities of tarsal and metatarsal bones. J Clin Densitom 2011;14:313 20. Available from: https://doi.org/10.1016/j. jocd.2011.05.006. [50] Commean PK, Ju T, Liu L, Sinacore DR, Hastings MK, Mueller MJ. Tarsal and metatarsal bone mineral density measurement using volumetric quantitative computed tomography. J Digit Imaging 2009;22:492 502. Available from: https://doi.org/10.1007/s10278-008-9118-z. [51] Chaplais E, Greene D, Hood A, Telfer S, du Toit V, Singh-Grewal D, et al. Reproducibility of a peripheral quantitative computed tomography scan protocol to measure the material properties of the second metatarsal. BMC Musculoskelet Disord 2014;15:242. Available from: https://doi. org/10.1186/1471-2474-15-242. [52] Bayod J, Becerro-de-Bengoa-Vallejo R, Losa-Iglesias ME, Doblare´ M. Mechanical stress redistribution in the calcaneus after autologous bone harvesting. J Biomech 2012;45:1219 26. Available from: https://doi.org/10.1016/j.jbiomech.2012.01.043. [53] Thomas VJ, Patil KM, Radhakrishnan S. Three-dimensional stress analysis for the mechanics of plantar ulcers in diabetic neuropathy. Med Biol Eng Comput 2004;42:230 5. [54] Wu L. Nonlinear finite element analysis for musculoskeletal biomechanics of medial and lateral plantar longitudinal arch of virtual Chinese human after plantar ligamentous structure failures. Clin Biomech (Bristol, Avon) 2007;22:221 9. Available from: https://doi.org/10.1016/j. clinbiomech.2006.09.009.

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[55] Telfer S, Erdemir A, Woodburn J, Cavanagh PR. Simplified vs geometrically accurate models of forefoot anatomy to predict plantar pressures: a finite element study. J Biomech 2016;49:289 94. Available from: https://doi.org/10.1016/j.jbiomech.2015.12.001. [56] Isvilanonda V, Dengler E, Iaquinto JM, Sangeorzan BJ, Ledoux WR. Finite element analysis of the foot: model validation and comparison between two common treatments of the clawed hallux deformity. Clin Biomech (Bristol, Avon) 2012;27:837 44. Available from: https://doi. org/10.1016/j.clinbiomech.2012.05.005. [57] Telfer S, Erdemir A, Woodburn J, Cavanagh PR. What has finite element analysis taught us about diabetic foot disease and its management? A systematic review. PLoS One 2014;9:e109994. Available from: https://doi.org/10.1371/journal.pone.0109994. [58] Fung A, Loundagin LL, Edwards WB. Experimental validation of finite element predicted bone strain in the human metatarsal. J Biomech 2017;60:22 9. Available from: https://doi.org/10.1016/j.jbiomech.2017.06.010. [59] van Eijnatten M, van Dijk R, Dobbe J, Streekstra G, Koivisto J, Wolff J. CT image segmentation methods for bone used in medical additive manufacturing. Med Eng Phys 2018;51:6 16. Available from: https://doi.org/10.1016/j.medengphy.2017.10.008. [60] Mueller MJ, Hastings M, Commean PK, Smith KE, Pilgram TK, Robertson D, et al. Forefoot structural predictors of plantar pressures during walking in people with diabetes and peripheral neuropathy. J Biomech 2003;36:1009 17. Available from: https://doi.org/10.1016/S0021-9290 (03)00078-2. [61] Ledoux WR, Rohr ES, Ching RP, Sangeorzan BJ. Effect of foot shape on the three-dimensional position of foot bones. J Orthop Res 2006;24:2176 86. Available from: https://doi.org/10.1002/jor.20262. [62] Anderson JG, Harrington R, Ching RP, Tencer A, Sangeorzan BJ. Alterations in talar morphology associated with adult flatfoot. Foot Ankle Int 1997;18:705 9. [63] Louie PK, Sangeorzan BJ, Fassbind MJ, Ledoux WR. Talonavicular joint coverage and bone morphology between different foot types. J Orthop Res 2014;32:958 66. Available from: https://doi.org/10.1002/jor.22612. [64] Peeters K, Natsakis T, Burg J, Spaepen P, Jonkers I, Dereymaeker G, et al. An in vitro approach to the evaluation of foot-ankle kinematics: performance evaluation of a custom-built gait simulator. Proc Inst Mech Eng Part H J Eng Med 2013;227:955 67. Available from: https://doi.org/ 10.1177/0954411913490455. [65] Schaefer KL, Sangeorzan BJ, Fassbind MJ, Ledoux WR. The comparative morphology of idiopathic ankle osteoarthritis. J Bone Jt Surg—Ser A 2012;94:e181. Available from: https://doi.org/10.2106/JBJS.L.00063 (1). [66] Ananthakrisnan D, Ching R, Tencer A, Hansen ST, Sangeorzan BJ. Subluxation of the talocalcaneal joint in adults who have symptomatic flatfoot. J Bone Jt Surg—Ser A 1999;81:1147 54. Available from: https://doi.org/10.2106/00004623-199908000-00010. [67] Tu¨mer N, Vuurberg G, Blankevoort L, Kerkhoffs GMMJ, Tuijthof GJM, Zadpoor AA. Typical shape differences in the subtalar joint bones between subjects with chronic ankle instability and controls. J Orthop Res 2019;37:1892 902. Available from: https://doi.org/10.1002/ jor.24336. [68] Telfer S, Kindig MW, Sangeorzan BJ, Ledoux WR. Metatarsal shape and foot type: a geometric morphometric analysis. J Biomech Eng 2017;139. Available from: https://doi.org/10.1115/1.4035077. [69] Moore ES, Kindig MW, McKearney DA, Telfer S, Sangeorzan BJ, Ledoux WR. Hind- and midfoot bone morphology varies with foot type and sex. J Orthop Res 2019;37:744 59. Available from: https://doi.org/10.1002/jor.24197. [70] Tu¨mer N, Blankevoort L, van de Giessen M, Terra MP, de Jong PA, Weinans H, et al. Bone shape difference between control and osteochondral defect groups of the ankle joint. Osteoarthr Cartil 2016;24:2108 15. Available from: https://doi.org/10.1016/j.joca.2016.07.015. [71] Alsayednoor J, Metcalf L, Rochester J, Dall’Ara E, McCloskey E, Lacroix D. Comparison of HR-pQCT- and microCT-based finite element models for the estimation of the mechanical properties of the calcaneus trabecular bone. Biomech Model Mechanobiol 2018;17:1715 30. Available from: https://doi.org/10.1007/s10237-018-1051-6. [72] Gao C, Chen Z, Cheng Y, Li J, Huang X, Wei L, et al. Comparative anatomy of the mouse and human ankle joint using micro-CT: utility of a mouse model to study human ankle sprains. Math Biosci Eng 2019;16:2959 72. Available from: https://doi.org/10.3934/mbe.2019146. [73] Mousavian A, Shakoor D, Hafezi-Nejad N, Haj-Mirzaian A, de Cesar Netto C, Orapin J, et al. Tibiofibular syndesmosis in asymptomatic ankles: initial kinematic analysis using four-dimensional CT. Clin Radiol 2019;74:571.e1 8. Available from: https://doi.org/10.1016/j. crad.2019.03.015.

Chapter 18

Weight-bearing Computed Tomography of the Foot and Ankle Franc¸ois Lintz1, Alessio Bernasconi2 and Cesar de Cesar Netto3 1

Ankle and Foot Surgery Center, Clinique de l’Union, Toulouse, Occitanie, France, 2Foot and Ankle Unit, Royal National Orthopaedic Hospital,

Stanmore, Greater London, United Kingdom, 3Department of Orthopedics and Rehabilitation, Carver College of Medicine, University of Iowa, Iowa City, Iowa, United States

Abstract Cone beam computed tomography (CT) was introduced in the 1990s, and it was initially utilized in the dental arena. However, since 2011, these machines have been used in the foot and ankle orthopedic setting. Several studies have demonstrated how cone beam computed tomography can provide weight-bearing images (hence, weight-bearing CT) with slices and reconstructed threedimensional (3D) models similar to traditional CT scans but acquired during physiological stance with markedly lower radiation exposure. This has allowed to overcome biases that are inherently related to two-dimensional radiographs and to test new 3D biometrics for the assessment of foot alignment. Fully automatic 3D measurements, bone segmentation, distance mapping, and modeling for custom-made implants represent the mainstays of the ongoing research on this new technology in foot and ankle surgery.

18.1

Introduction

When Wilhelm Conrad Roentgen discovered radiography in 1895, he reported “I’ve discovered something interesting, but I don’t know if my observations are correct.” Of course, two-dimensional (2D) radiography revolutionized medicine during the next century, enabling at the time an unprecedented in vivo exploration of the human body. Nevertheless, Roentgen was probably right when doubting himself in that radiography has always generated biased and somehow distorted projections of the body, not providing an accurate representation of reality itself [1 4]. With the advent of computed tomography (CT) in the 1970s, the skeleton was revealed in 3D [5]. Many medical and surgical disciplines utilized these developments in the imaging field, with improved understanding of pathophysiology and increased diagnostic power. Orthopedics, in particular foot and ankle surgery, was not an exception, with radiographs and CT scanners soon becoming essential parts of the diagnostic and therapeutic pathway. While X-rays allow the foot and ankle to be studied in a standing (weight bearing) condition with a relatively low spatial resolution and superimposition of structures, a CT scan is generally requested to obtain more detail on the bones and joints. This is usually performed with the patient supine and involves a high radiation exposure compared to standard X-rays. Most clinical CT scanners operate in a helical mode where both the X-ray tube and the detector rotate uninterruptedly while the patient is lying on a table which translates through a gantry. Considering the table movement and gantry characteristics (height and diameter of about 2 meters and weight around 2000 kg), it would be very difficult to create an upright machine with vertical movement either of the table (at risk of artifacts coming from patient movements) or of the gantry (heavy to mobilize). Although the absence of gravity makes this imaging data evidently unreliable in the assessment of “true” foot bone positions (where true means in physiological standing position), over time doctors’ brains have been trained to integrate the two imaging modalities and use them to gather useful data about the patient’s anatomy and pathophysiology. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00012-3 © 2023 Elsevier Inc. All rights reserved.

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As found in the scientific literature, several attempts have been made to overcome these limitations. Several studies have documented the use of CT scans of the foot and ankle under “simulated” weight-bearing conditions using different types of devices, although due to difficulties in fully loading the body in a clinical CT scanner these have “partially” reproduced the loading condition and lack active muscle activation that plays a key role in determining the relative position of the bones [6]. An important step forward has been then represented by the introduction of cone beam weight-bearing CT (WBCT), which offers a reasonable solution for obtaining three-dimensional (3D) weight-bearing skeleton images that are important to our understanding of deformities and degenerative pathologies affecting the lower limbs [1]. In practice, cone beam WBCT combines the advantages of 3D imaging and high spatial resolution found in CT, with the ability to stand during scanning and a low radiation dose comparable to a 2D X-ray, allowing doctors to accurately measure the relative positions of the bones under full weight bearing. Of course, this does not come without challenges for the clinical and scientific communities, pertaining to the fact that existing tools, which applied to 2D radiography, are not adapted to the new 3D environment. Indeed, it is more through the disappearance of 2D projection biases that 3D WBCT may potentially transform lower limb and especially foot and ankle imaging. WBCT may also increase the accessibility of personalized risk assessment and personalized surgery due to its lower radiation dose.

18.2

Biases of conventional radiography

Undoubtedly, the main advantage of cone beam WBCT is to produce an accurate 3D model of the weight-bearing foot, thereby overcoming most of the biases related to conventional 2D radiography [6]. It is, therefore, essential to understand what these biases are. First, perspective biases among which are rotation bias and the fan effect [2,4]. The former depends on the fact that the incidence angle of the X-ray beam on the body varies, with shapes and angles on the film changing accordingly. The latter is a consequence of the distance from the X-ray source to the object, since projected lengths will increase compared to the real length. Second, there is an operator-related bias. This concerns the technical aspects of setting the radiographic apparatus around the patient and the placement of the patient. The positioning of the foot opposite to the X-ray source, the height of the source and its distance are impossible to reproduce exactly from one setup to the other. In practice, patients are generally radiographed at least twice to obtain an anterior-posterior and a lateral view, which is the standard combination in orthopedic imaging. Conversely, the use of software allows the user to manipulate 3D WBCT images and to obtain any number of views from which the observer may define and review according to their interest. In the future, the introduction of dedicated algorithms may enable standard views from 3D data to be harmonized, making them comparable between patients. Third, we must consider superimposition bias, related to the projection of a 3D structure in a 2D plane, where multiple planes are piled up into a single plane [7] (Fig. 18.1). This results in areas where contours cannot be distinguished, where edges can be superimposed, and gives images which require a certain amount of guesswork or experience to decipher. This is what the orthopedic foot and ankle surgeons’ brains have been trained to do for decades. All these biases result in the fact that no radiograph is reliably reproducible unless it can be ensured that the radiographic setup can be reinstalled in exactly the same way twice for the same patient, which is not possible. However, despite this fact seeming intuitively true, it has been largely ignored by the literature which has systematically focused

FIGURE 18.1 Example of standard radiograph of the left foot with anterior/posterior (left) and medial/lateral (right) views to demonstrate the superimposition bias related to 2D imaging. The borders of midfoot bones (middle and lateral cuneiform, navicular, cuboid) and the base of the second to fifth metatarsal bones are difficult to delineate since they are superimposed. As such, in case of suspected bony lesion on these areas, there is a very low threshold to request secondary imaging like computed tomography.

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on intra and inter-observer reliability, even developing new measurements to improve the latter [8 11]. Hindfoot alignment exemplifies this concept through the great number of techniques described: Saltzman view, Meary view, long axial view, and hindfoot alignment view only to cite the most popular of them [8 11]. All these techniques focus on being reproducible rather than accurate; the measurement itself never corresponds to the “real” alignment of the hindfoot because the radiograph itself is only a projection of the real foot and ankle complex. The result is that a measurement technique may be perfect, but the result is still biased since applied to an altered projection of the real angle. A recent study showed that when the foot is rotated 30 degrees away from the neutral position (including X-rays aligned with the second metatarsals), the hindfoot alignment measurement changes by 1% for each supplementary degree of rotation [4]. In terms of consequence for patients, another study showed that the mean variability of radiographic measurements away from the “real angle” was approximately 20% [12]. So it is possible to measure a varus hindfoot position in a case where the “true” or 3D position is valgus. Similarly, Willauer et al. have demonstrated how X-ray source misalignment up to 25 and 30 degrees on the transverse and sagittal planes, respectively, may lead to errors in measuring common angles like the calcaneal pitch and the talonavicular coverage angle [2]. This has consequences in terms of surgical planning of corrective osteotomies in complex hindfoot deformities.

18.3

Technical aspects

Cone beam CT is a relatively recent technology, first published in 1998 by Mozzo et al. [13]. It allows the reconstruction of 3D models from the information contained in stacked 2D transverse slices of the anatomy, much like a conventional CT scan [14]. In the case of cone beam CT, however, a cone beam (instead of a fan beam as seen in traditional CT) is projected through the anatomy (Fig. 18.2). The result is called a sinogram. This is the continuous projection of the anatomy on the target (a standard flat panel detector), which faces the X-ray source on the other side of the patient’s foot and ankle. To decipher this image back to a 3D volume, mathematical algorithms based on the Radon and the Fourier transforms are necessary. Fourier tells us how to distinguish multiple signals piled up together. Radon tells us how to calculate or back-project the coordinates of each pixel and therefore reconstruct the whole volume, slice after slice [14]. The stability of the 3D matrix in which the foot and ankle complex 3D model is acquired remains contained within the machine’s gantry. Any variation, which could be due to mechanical loosening of its parts or aging of the machine, is regularly controlled through industry-standard QA (Quality Assessment) procedures, as is the case for existing machines. This consists of scanning Plexiglas templates containing materials of known dimensions and densities to detect any variation compared to the known values. In cone beam CT, the X-ray source only performs a single revolution around the anatomy, compared to a conventional CT which rotates many times during a scan (Fig. 18.2). The result is that a much lower radiation dose is required to obtain the image volume data (typical effective radiation dose is 0.001 mSv for single-exposure X-ray of the foot, 0.07 mSv for a conventional CT of the ankle, and 0.01 0.03 mSv for a cone beam CT of the foot/ankle) [6], which makes an important difference in the clinical setting. For example, considering a standard knee or foot and ankle assessment prior to a reconstructive surgical procedure, it typically requires some 3D nonweight-bearing information (provided by conventional CT) which needs to be coupled with 2D weight-bearing information (6 8 radiographs, usually including bilateral anteroposterior views, lateral, and complementary views). Merging all this information happens in

FIGURE 18.2 Illustration of a fan-shaped beam (left image) and cone-shaped beam (right image). While the first one requires multiple rotations to capture the images, the second one needs one single rotation with significant reduction in radiation dose. Courtesy Curvebeam.

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the surgeons’ brains and requires experience, and some degree of planning, based on the biased 2D measurements. On the contrary, the radiological cost of WBCT is only of 0.01 mSv [for a small field of view (FOV), typically 15 cm in diameter scanning a single foot] to 0.06 mSv for a large 30 cm FOV scanning bilateral feet. This corresponds to approximately 1% of the yearly background radiation dose in the United States (3 mSv/y) and is equivalent to 10 conventional radiographs, but is able to produce a precise virtual 3D weight-bearing model of the anatomy of the two lower limbs [6] (Fig. 18.3). Understanding this model, however, requires moving to the interpretation of 3D imaging as opposed to 2D, which is one of the challenges faced when adopting WBCT technology. In the 3D environment that contains the foot and ankle virtual model of the patient’s anatomy (i.e., the FOV), what used to be a flat picture becomes a volume and pixels become voxels. The grayscale of each voxel is defined by Hounsfield’s attenuation coefficient (HU for Hounsfield Unit), which corresponds to the density of the traversed anatomical tissues compared to air and water (where the value for air is 21000 and 0 for water). Furthermore, where landmark positions were previously defined in 2D by their relative position to other landmarks through distances and angles (generating the mentioned biased measurements), it is now possible to define the position of each voxel relative to an orthogonal referential. Each voxel is now defined by its X, Y, and Z coordinates and its HU. In terms of size, the acquisition volume within a typical bilateral WBCT contains 1,000,000 voxels each approximately 0.3 3 0.3 3 0.3 mm in size. A typical foot within this space would be represented by approximately 200,000 voxels. Acquisition time for such a volume is typically under a minute. Because interacting with a 3D virtual model of the foot and ankle is completely different to interacting with a flat image, a different set of tools is required. Depending on the software provider, the presentation of the 3D model may vary but generally, an MPR (multi planar reconstruction) screen is used, divided in 4 parts, containing one 3D visual (3D rendering) and the 3 spatial 2D planes (coronal, transverse, and sagittal) allowing the user to scroll through the image slices in different planes or navigate the 3D model and make measurements (Fig. 18.4). Along with this “modern” presentation, most software packages allow users to also view digitally reconstructed radiographs (DRRs). DRRs are artificial flat, 2D conventional radiographs which are reconstructed from the information contained in the 3D volume. In algorithmic terms, this is achieved by increasing the width of a chosen slice (a single “slice” of the volume) in a chosen plane (Fig. 18.5). This postprocessing option has two main reasons for existing. First, the current generation of specialists has been trained using 2D radiographs so it will take some time before the newer 3D technology will be accepted. Second, in clinical practice it may be useful at times to get an “overall” image like a snapshot rather than single CT slices. For that purpose, DRRs are interesting since they can be obtained without any extra radiation exposure to the patient (as opposed to the older protocols that combined conventional CT and radiographs).

FIGURE 18.3 Example of a 3D rendered bilateral cone beam CT scan. Skin rendering (left) and bone rendering (right).

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FIGURE 18.4 MPR (Multi Planar Reconstruction) window from a bilateral WBCT scan on Cubevue software (Curvebeam LLC, Hatfield PA). In the upper left a 3D reconstruction of both feet is shown seen. The other three quadrants show the three planes of the volume.

FIGURE 18.5 Example of a Digitally Reconstructed Radiograph (DDR) obtained from cone beam WBCT images to reproduce a Saltzman view and assess the hindfoot alignment in a traditional t2D way.

Clinical use of cone beam CT was first implemented in the dental arena in the late 1990s. This was followed by a period of approximately 10 years in which the technology gained popularity in this field, until the time of writing where cone beam CT has gained widespread popularity and in many cases has replaced conventional 2D panoramic radiographs. This natural evolution of its use in the dental field has been typical of any new technology, with an initial period of rapid adoption by innovators who believed in its potential despite initially high costs that are later reduced by increases in production efficiency related to mass adoption.

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In orthopedics, the use of 3D images has already enabled the use of patient-specific computerized solutions (Patient Specific Instrumentation) such as custom surgical guides and even custom implants [15,16]. This has been more common in the field of knee surgery, but recently also in ankle arthroplasty. Potentially, the use of cone beam technology may allow the surgical team to obtain weight-bearing images that can be used to improve the accuracy of cutting guides and the design of the implant. In terms of costs, a typical WBCT machine is still relatively expensive, around $150,000 250,000. Although an accurate cost analysis has never been performed to date, the use of WBCT with relatively quick scanning times (under a minute) and the possibility to obtain in-office 3D bilateral imaging may help reduce waiting lists in radiology and free up conventional CT slots for other patients as well as reduce costs. Furthermore, even if conventional CT was as fast as cone beam CT, the complete process of getting conventional radiographs for weight-bearing measurements and then adding a fan beam CT to obtain 3D images would be on average a lot more time-consuming, as was demonstrated in a recent study by Richter et al. [17]. Summary data from one of the first clinics to implement WBCT recently reported savings of 16,000 hours/year of unnecessary imaging time and 6000 mSV of radiation (the equivalent of 750 years of US daily radiation doses). This was explained by the drastic cut, during the same period on the consumption of standard radiographs (283%) and traditional CT scans (296%) [17].

18.4

Indications

Researchers and clinicians began publishing on WBCT around 2011, although partial weight-bearing was an active area of research before this, and interest in the field has been steadily growing since then. The main subjects of interest have been: flat foot [18 21], the subtalar joint [22,23], ankle fractures [24], syndesmotic injuries [25 28], ankle instability [29], and hallux rigidus [30]. These recent papers have focused on identifying or confirming anatomical modifications associated with pathological conditions in the foot and ankle, and initial findings are described next. Concerning flat feet and the subtalar joint, flat foot measurements analogous to conventional radiologic parameters may be obtained using WBCT and have better detection of severity [18 20]; in an uninjured population, the fifth metatarsal demonstrates plantarflexion relative to the first metatarsal in patients with flat feet relative to controls [21]; patients with flat foot deformity were found to have more innate valgus in their talar shape and in their subtalar alignment [22]; subtalar joint orientation may be a risk factor for the development of ankle joint osteoarthritis [22,23]. Concerning the tibiofibular mortise, there is internal rotation of the talus in the varus osteoarthritic ankle, increasing with severity [26]; weight-bearing rotation of the talus within the normal mortise is around 10 degrees, fibular posterior translation is 1.5 mm and external rotation 3 degrees, and comparison with the contralateral side seems to be more reliable than with the population norm [27]. A recent paper reported no significant changes of the distal tibiofibular joint measurements in WBCT compared to a nonweight bearing setup in asymptomatic uninjured ankles but the medial clear space of the tibiotalar joint was increased in the latter, likely due to the anterior shift of the talus in a nonweightbearing configuration because of the greater width of the dome larger anteriorly compared to posteriorly [28]. A greater risk of chronic lateral ankle instability has been proven and quantified in patients with varus hindfoot malalignment [29]. With regards to the first ray, hallux rigidus patients have metatarsus primus elevatus, increasing with the severity [30]. In our experience, indications for WBCT have mostly been for conditions where the combination of 3D imaging during weight bearing seemed appealing. However, we suggest that in the future, the whole range of foot and ankle conditions may be considered in a more straightforward way through this new technology. In hallux valgus for example, the role of coronal rotational position of the first metatarsal and sesamoids needs to be taken into account when planning a surgery correction [31 34]. It seems logical to anticipate that in the future the understanding, diagnosis, and treatment of forefoot conditions may also benefit from WBCT.

18.5

3D biometrics

It is in the area of preoperative planning that WBCT is the most promising, allowing personalized assessment, devoid of conventional imaging biases [6,35,36]. In practice, the possibility of obtaining detailed images during standing to reproduce the physiological position of the limb in the space might enable measurements of the anatomy to be taken under the effect of load and subsequently allow the surgeon to plan with greater accuracy. Until recently, measurements used for orthopedic preoperative planning have been predominantly based on 2D radiographs. Due to the subjectivity of clinical assessment, distances and angles have represented a more reliable way to “standardize” the description of deformities, allowing patients to be compared, and to verify changes before and after treatment. After the introduction of WBCT, a few authors have showed how traditional 2D measurements may be used on 3D images [18 28,30].

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These measurements are by definition still subject to rotation bias, however, since the 3D volume can be manipulated after having been acquired, the rotational position of the foot model in which the measurement is made can be determined with more precision or modified without having to obtain further X-rays [6]. Three methods have been advocated in literature to measure angles and distances using WBCT datasets: (1) measurements are made in DRR mode [35,37] (Fig. 18.5); (2) measurements are made in a single slice [18 22]; or (3) measurements are projected onto the 3D rendering volume [38]. In the first and third cases, this is the same as plain radiographs, with the same biases. In the second case, the difficulty lies in defining the slice in which the measurement is made. To increase reproducibility this can be based on surface anatomical markers which unfortunately have their own variability. In obtaining true 3D measurements, allowing markers to be placed in the three planes simultaneously, solutions exist which require third party software providers, but are often costly and time-consuming because they require prior segmentation of the bony architecture of the foot and ankle complex [38]. One possible solution to this problem may be 3D biometrics, a concept recently published by Lintz et al. [39]. 3D biometrics are anatomical measurements defined by three key features: computerized, semiautomatic, and volumetric (i.e., based on at least four points). In this case, a new measurement for hindfoot alignment dubbed the foot ankle offset (FAO) is obtained using a specific software (Talas, Curvebeam, LLC, Hatfield PA). Traditional measurements for hindfoot alignment usually define the position of the projection of the calcaneus through its position relative to the anatomical axis of the tibia. The proposed 3D biometric system analyzes the position of the center of the ankle joint, relative to the three actual (as opposed to projected) weight-bearing points of the foot (the first and second metatarsal heads and the calcaneus). This is inspired by previous work by Saltzman et al. [8], Arunakul et al. [40] and Lintz et al. [12] who had respectively proposed replacing the tibial-calcaneal angle by an offset [8], and to define the position of the hindfoot relative to the forefoot rather than the tibia [12,40]. The advantages are that, contrary to an angle, an offset is directly related to the lever arm, thus, to the torque generated by the distance between where body weight and ground reaction force respectively are applied on the foot and ankle complex [12,39]. In the aforementioned semiautomatic software, the investigator indicates where the four points are situated using an MPR window setting (Fig. 18.6), thereafter the software automatically calculates the value of the FAO [29,39,41 43].

X

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8.41

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FIGURE 18.6 3D, semiautomatic biometric measurement of hindfoot alignment using the Foot Ankle Offset (FAO), performed with a dedicated software (TALAS) on a normally-aligned (upper image, FAO: 0.4%), varus (lower left image, FAO: 215.9%) and valgus (lower right image, FAO: 8.4%) hindfoot. Foot and Ankle Offset is described in %, where the absolute value of the offset is normalized relative to the length of the foot. A normal value has been assessed at around 1% 2%, with lower values representing varus alignment and higher values valgus.

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Advantages and limitations of weight-bearing computed tomography

Summing up what has been discussed, the main advantages of weight-bearing CT technology may be listed as follows: G

G

G

G

possible to obtain images with detail comparable to traditional CT but in a stance position, reproducing the position of anatomical structures under load [1,6]; overall radiation exposure markedly less than CT and similar to a complete set of foot and ankle radiographs (including dorsoplantar and lateral views of the feet, anteroposterior view of the ankle, and specific calcaneal views for hindfoot alignment) [6,35,36]; limited physical amount of space required to host the machine [a typical machine would fit in a 3 m 3 3 m room (PedCat, Curvebeam LLC, Hatfield, PA; Verity, Planmed, Helsinki, Finland; Onsight 3D Extremity System, Carestream Inc, Rochester, NY)] calculation of 3D biometrics through a dedicated software to overcome biases inherently related to 2D measurements [29,39,41 43]. On the other hand, some limitations of WBCT must also be considered, namely:

G

G

G

G

G

A high initial cost (although lower than most conventional medical CTs). This may be compensated by time savings, however to date no specific cost-analysis has been performed. Its availability is still confined to a large center, likely due to its recent introduction on the market in 2014, thus limiting the amount of research on daily clinical practice. While the FAO represents a 3D tool to investigate foot alignment with excellent reliability [39,41 43], it remains the only available 3D measurement applicable on 3D images, and further research is warranted on its use in clinical routine; many authors, to date, are still using classical radiographic angles and distances on WBCT images [18 20,22], probably not taking advantage of the 3D nature of data. A precise and automatic segmentation of the foot will likely allow researchers and clinicians to better investigate bone position [38], rotation and axes, however to the best of authors’ knowledge so far software are still unable to accurately recognize bone borders, perhaps due to severe degenerative changes. Regarding the quality of cone beam imaging, the literature is still controversial. Some authors have documented good quality images from weight-bearing scanners with high reliability in common measurements [1,18], while according to Lechuga and Weidlich fan beam technology would produce a better soft tissue differentiation than cone beam [44]. Further studies will shed more light on this area.

18.7

Future areas of research

One of the major aims of ongoing research in WBCT and other medical imaging technologies is to obtain reliable fully automatic measurement systems. These have the potential to save large amounts of clinical time, avoid the need to manually define landmarks [39], and potentially further increasing reproducibility. Another objective in the WBCT area is to achieve automatic segmentation of the foot and ankle bones. Currently, this is not possible, but efforts are being made by the industry and the academic world to develop algorithms that can perform this task. In theory, once the anatomy is segmented, the computer will know which bone is which and determine axes and relative orientations. When considering complex foot deformity like severe flat foot or neurological pes cavovarus, it is understandable how this type of 3D approach on weight-bearing images may help to analyze the morphotype, accurately plan the correction, and verify the result achieved. A further area of development is distance mapping [38] (Fig. 18.7). This technique studies the joint surface interaction in a noninvasive fashion by using 3D bone models gathered from 3D images. A recent study by Siegler has demonstrated how distance mapping provides detailed data about ankle and subtalar joints during normal movements of the joints [38]. If confirmed by further clinical studies, this may represent an important tool in the understanding of joint degeneration, helping in diagnosis and treatment planning. Finally, 3D models from WBCT images might be used to design custom-made instruments and implants. Although this is already occurring using traditional CT reconstruction [15,16], the weight-bearing setting may enhance the accuracy of final products while reducing radiation exposure for the patient. Further studies are warranted to verify this hypothesis. Ongoing improvements are being made to the software used by WBCT and other medical imaging scanners to help reduce noise and provide better image quality. New technologies such as photon counting are being explored and may

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FIGURE 18.7 Example of distance mapping technology from WBCT images. The process to obtain surface-to-surface joint interaction maps (indicated with a range of colors) is still dependent on human intervention for segmentation and sequence. The maps are related to pressure distribution. Research is ongoing to make the process fully automatic.

allow different kinds of soft tissues to be distinguished. These technologies may require the application of artificial intelligence to medical imaging data, an area that is currently being investigated intensively by the academic world and the industry.

18.8

Conclusion

Cone beam CT technology can be used to obtain accurate 3D images of the foot and ankle complex during weight bearing, with low radiation and a similar spatial resolution when compared to traditional CT. Multiple studies on various foot and ankle conditions have shown that both classical 2D measurements and new 3D biometric tools have excellent reliability on weight-bearing CT images. Studies are warranted to confirm its cost-effectiveness and high quality of images. Further research is required to validate new 3D measurements, to obtain fully automatic bone segmentation, to apply distance mapping techniques to clinical practice, and to test the validity of WBCT 3D reconstruction as a model for custom-made implants.

Acknowledgments The authors are grateful to the International Weight Bearing CT Society for their support during the editing of this chapter, in particular Professor Martinus Richter, Dr Alexej Barg, and Dr Arne Burssens.

Conflict of interest statement Franc¸ois Lintz and Cesar de Cesar Netto declare paid consultancy from Curvebeam. Franc¸ois Lintz has a patent for TALAS software.

References [1] Richter M, Seidl B, Zech S, Hahn S. PedCAT for 3D-imaging in standing position allows for more accurate bone position (angle) measurement than radiographs or CT. Foot Ankle Surg 2014;20(3):201 7. [2] Willauer P, Sangeorzan BJ, Whittaker EC, Shofer JB, Ledoux WR. The sensitivity of standard radiographic foot measures to misalignment. Foot Ankle Int 2014;35(12):1334 40. [3] Barg A, Amendola RL, Henninger HB, Kapron AL, Saltzman CL, Anderson AE. Influence of ankle position and radiographic projection angle on measurement of supramalleolar alignment on the anteroposterior and hindfoot alignment views. Foot Ankle Int 2015;36(11):1352 61. [4] Baverel L, Brilhault J, Odri G, Boissard M, Lintz F. Influence of lower limb rotation on hindfoot alignment using a conventional twodimensional radiographic technique. Foot Ankle Surg 2017;23(1):44 9. [5] Ambrose J, Hounsfield G. Computerized transverse axial tomography. Brit J Radiol 1973;46:148 9. [6] Barg A, Bailey T, Richter M, de Cesar Netto C, Lintz F, Burssens A, et al. Weightbearing computed tomography of the foot and ankle: emerging technology topical review. Foot Ankle Int 2018;39(3):376 86. [7] Perlepe V, Omoumi P, Larbi A, Putineanu D, Dubuc JE, Schubert T, et al. Can we assess healing of surgically treated long bone fractures on radiograph? Diagn Interv Imaging 2018;99(6):381 6.

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[8] Saltzman CL, el-Khoury GY. The hindfoot alignment view. Foot Ankle Int 1995;16(9):572 6. [9] Reilingh ML, Beimers L, Tuijthof GJ, Stufkens SA, Maas M, van Dijk CN. Measuring hindfoot alignment radiographically: the long axial view is more reliable than the hindfoot alignment view. Skelet Radiol 2010;39(11):1103 8. [10] Williamson ER, Chan JY, Burket JC, Deland JT, Ellis SJ. New radiographic parameter assessing hindfoot alignment in stage II adult-acquired flatfoot deformity. Foot Ankle Int 2015;36(4):417 23. [11] Dagneaux L, Moroney P, Maestro M. Reliability of hindfoot alignment measurements from standard radiographs using the methods of Meary and Saltzman. Foot Ankle Surg 2017; Nov 11. pii: S1268 7731(17)31331-0. [12] Lintz F, Barton T, Millet M, Harries WJ, Hepple S, Winson IG. Ground reaction force calcaneal offset: a new measurement of hindfoot alignment. Foot Ankle Surg 2012;18(1):9 14. [13] Mozzo P, Procacci C, Tacconi A, Martini PT, Andreis IA. A new volumetric CT machine for dental imaging based on the cone-beam technique: preliminary results. Eur Radiol 1998;8(9):1558 64. [14] Scarfe WC, Farman AG. What is cone-beam CT and how does it work? Dent Clin North Am 2008;52(4):707 30. [15] Berlet GC, Penner MJ, Lancianese S, Stemniski PM, Obert RM. Total ankle arthroplasty accuracy and reproducibility using preoperative CT scan-derived, patient-specific guides. Foot Ankle Int 2014;35(7):665 76. [16] Wagener J, Gross CE, Schweizer C, Lang TH, Hintermann B. Custom-made total ankle arthroplasty for the salvage of major talar bone loss. Bone Jt J 2017;99-B(2):231 6. [17] Richter M, Lintz F, de Cesar Netto C, Barg A, Burssens A. Results of more than 11,000 scans with weightbearing CT—impact on costs, radiation exposure, and procedure time. Foot Ankle Surg 2019; Jun 18. pii: S1268 7731(19)30096-7. [18] de Cesar Netto C, Schon LC, Thawait GK, et al. Flexible adult acquired flatfoot deformity. J Bone Jt Surg 2017;99(18):e98. [19] de Cesar Netto C, Shakoor D, Dein EJ, Zhang H, Thawait GK, Richter M, et al. Influence of investigator experience on reliability of adult acquired flatfoot deformity measurements using weightbearing computed tomography. Foot Ankle Surg 2018; Mar 12. pii: S1268 7731(18) 30058-4. [20] de Cesar Netto C, Shakoor D, Roberts L, Chinanuvathana A, Mousavian A, Lintz F, et al. Hindfoot alignment of adult acquired flatfoot deformity: a comparison of clinical assessment and weightbearing cone beam CT examinations. Foot Ankle Surg 2018; Nov 5. pii: S1268 7731(18) 30301-1. [21] Yoshioka N, Ikoma K, Kido M, Imai K, Maki M, Arai Y, et al. Weight-bearing three-dimensional computed tomography analysis of the forefoot in patients with flatfoot deformity. J Orthop Sci 2016;21(2):154 8. [22] Probasco W, Haleem AM, Yu J, Sangeorzan BJ, Deland JT, Ellis SJ. Assessment of coronal plane subtalar joint alignment in peritalar subluxation via weight-bearing multiplanar imaging. Foot Ankle Int 2015;36(3):302 9. [23] Kra¨henbu¨hl N, Tschuck M, Bolliger L, Hintermann B, Knupp M. Orientation of the subtalar joint: measurement and reliability using weightbearing CT scans. Foot Ankle Int 2016;37(1):109 14. [24] Lawlor MC, Kluczynski MA, Marzo JM. Weight-bearing cone-beam CT scan assessment of stability of supination external rotation ankle fractures in a cadaver model. Foot Ankle Int 2018;39(7):850 7. [25] Osgood GM, Shakoor D, Orapin J, Qin J, Khodarahmi I, Thawait GK, et al. Reliability of distal tibio-fibular syndesmotic instability measurements using weightbearing and non-weightbearing cone-beam CT. Foot Ankle Surg 2018; Oct 31. [26] Lepoja¨rvi S, Niinima¨ki J, Pakarinen H, Koskela L, Leskela¨ HV. Rotational dynamics of the talus in a normal tibiotalar joint as shown by weight-bearing computed tomography. J Bone Jt Surg [Am] 2016;98:568 75. [27] Lepoja¨rvi S, Niinima¨ki J, Pakarinen H, Leskela¨ HV. Rotational dynamics of the normal distal tibiofibular joint with weight-bearing computed tomography. Foot Ankle Int 2016;37:627 35. [28] Shakoor D, Osgood GM, Brehler M, Zbijewski WB, de Cesar Netto C, Shafiq B, et al. Cone-beam CT measurements of distal tibio-fibular syndesmosis in asymptomatic uninjured ankles: does weight-bearing matter? Skelet Radiol 2019;48(4):583 94. [29] Lintz F, Bernasconi A, Baschet L, Fernando C, Mehdi N, Weight Bearing CT International Study Group, et al. Relationship between chronic lateral ankle instability and hindfoot varus using weight-bearing cone beam computed tomography. Foot Ankle Int 2019; Jun 28:1071100719858309. [30] Cheung ZB, Myerson MS, Tracey J, Vulcano E. Weightbearing CT scan assessment of foot alignment in patients with hallux rigidus. Foot Ankle Int 2018;39(1):67 74. [31] lamo-Espinosa JM, Flo´rez B, Villas C, et al. The relationship between the sesamoid complex and the first metatarsal after hallux valgus surgery without lateral soft- tissue release: a prospective study. J Foot Ankle Surg 2015;54:1111 15. [32] Geng X, Zhang C, Ma X, et al. Lateral sesamoid position relative to the second metatarsal in feet with and without hallux valgus: a prospective study. J Foot Ankle Surg 2016;55:136 9. [33] Chen JY, Rikhraj K, Gatot C, Lee JY, Singh Rikhraj I. Tibial sesamoid position influence on functional outcome and satisfaction after hallux valgus surgery. Foot Ankle Int 2016;37:1178 82. [34] Welck MJ, Singh D, Cullen N, Goldberg A. Evaluation of the 1st metatarso-sesamoid joint using standing CT—the Stanmore classification. Foot Ankle Surg 2018;24(4):314 19. [35] Lintz F, de Cesar Netto C, Barg A, Burssens A, Richter M, Weight Bearing CT International Study Group. Weight-bearing cone beam CT scans in the foot and ankle. EFORT Open Rev 2018;3(5):278 86. [36] Godoy-Santos Al, Netto Cesar C, Weight-Bearing Ct International Study Group. Weight-bearing computed tomography of the foot and ankle: an update and future directions. Acta Ortop Bras 2018;26(2):135 9.

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[37] Moore CS, Wood TJ, Saunderson JR, Beavis AW. A method to incorporate the effect of beam quality on image noise in a digitally reconstructed radiograph (DRR) based computer simulation for optimisation of digital radiography. Phys Med Biol 2017;62(18):7379 93. [38] Siegler S, Konow T, Belvedere C, Ensini A, Kulkarni R, Leardini A. Analysis of surface-to-surface distance mapping during three-dimensional motion at the ankle and subtalar joints. J Biomech 2018;76:204 11 Jul 25. [39] Lintz F, Welck M, Bernasconi A, Thornton J, Cullen NP, Singh D, et al. 3D biometrics for hindfoot alignment using weightbearing CT. Foot Ankle Int 2017;38(6):684 9. [40] Arunakul M, Amendola A, Gao Y, Goetz JE, Femino JE, Phisitkul P. Tripod index: a new radiographic parameter assessing foot alignment. Foot Ankle Int 2013;34(10):1411 20. [41] Zhang JZ, Lintz F, Bernasconi A, Weight Bearing CT International Study Group, Zhang S. 3D biometrics for hindfoot alignment using weightbearing computed tomography. Foot Ankle Int 2019;. Available from: https://doi.org/10.1177/1071100719835492 Mar 10:1071100719835492. [42] de Cesar Netto C, Bernasconi A, Roberts L, Pontin PA, Lintz F, Saito GH, et al. Foot alignment in symptomatic national basketball association players using weightbearing cone beam computed tomography. Orthop J Sports Med 2019;7(2). Available from: https://doi.org/10.1177/ 2325967119826081 2325967119826081. [43] Bernasconi A, Cooper L, Lyle S, Patel S, Cullen N, Singh D, et al. Intraobserver and interobserver reliability of cone beam weightbearing semiautomatic three-dimensional measurements in symptomatic pes cavovarus. Foot Ankle Surg 2019; Jul 26. pii: S1268-7731(19)30110-9. [44] Lechuga L, Weidlich GA. Cone beam CT vs. fan beam CT: a comparison of image quality and dose delivered between two differing CT imaging modalities. Cureus. 2016;8(9):e778.

Further reading Richter M, Frink M, Zech S, Vanin N, Geerling J, Droste P, et al. Intraoperative pedography: a validated method for static intraoperative biomechanical assessment. Foot Ankle Int 2006;27(10):833 42. Richter M, Lintz F, Zech S, Meissner SA. Combination of PedCAT weightbearing CT with pedography assessment of the relationship between anatomy-based foot center and force/pressure-based center of gravity. Foot Ankle Int 2018;39(3):361 8. Richter M, Zech S, Hahn S, Naef I, Merschin D. Combination of pedCATs for 3D imaging in standing position with pedography shows no statistical correlation of bone position with force/pressure distribution. J Foot Ankle Surg 2016;55(2):240 6. Richter M, Zech S. Leonard J. Goldner Award 2009. Intraoperative pedobarography leads to improved outcome scores: a Level I study. Foot Ankle Int 2009;30(11):1029 36. Richter M., Lintz F., Netto C.C., Barg A., Burssens A., Ellis S. Weight Bearing Cone Beam Computed Tomography (WBCT) in the Foot and Ankle. A Scientific, Technical and Clinical Guide. Springer.

Chapter 19

Magnetic Resonance Imaging of the Foot and Ankle Tim Finkenstaedt1,2,3, Palanan Siriwanarangsun1,4 and Christine B. Chung1,5 1

UC San Diego Department of Radiology, La Jolla, CA, United States, 2Institute of Diagnostic and Interventional Radiology, University Hospital

Zurich, Zurich, Switzerland, 3Institute of Radiology and Nuclear Medicine, Kantonsspital Winterthur, Winterthur, Switzerland, 4Department of Radiology, Siriraj Hospital, Mahidol University, Bangkok Noi, Bangkok, Thailand, 5VA San Diego Healthcare System, Department of Radiology, San Diego, CA, United States

Abstract Magnetic resonance imaging (MRI) can provide detailed information regarding the anatomy of the foot and ankle in health and disease. The physics and interpretation of MRI are complicated, and here we provide helpful concise tenets to facilitate understanding of the fundamental principles of MRI physics and interpretation in clinical routine. The advantages and disadvantages of MRI versus computed tomography (CT) for the musculoskeletal system are briefly discussed. Furthermore, the basic MRI appearance of normal and pathologic musculoskeletal tissue is demonstrated and a tailored MRI protocol for the foot and ankle region is provided. The main emphasis is on the anatomical detail and common pathology of the respective areas (ankle, hindfoot, midfoot, and forefoot).

19.1

Introduction

The working principle of magnetic resonance imaging (MRI) is briefly explained. MRI is based on the electromagnetic activity of atomic nuclei. In routine medical imaging, hydrogen (H1) nuclei are used most often because of their abundance in the body. Only nuclei that have a net positive spin are considered MR-active nuclei [1]. With even numbered nuclei, electrons and protons result in neutral charge and no ability to leverage the proton energy state. By inducing a magnetic field around the patient, these nuclei can be used to generate a signal (Fig. 19.1). Different tissues have different T1, T2, and T2-star values (unit: milliseconds) and, thereby, different “weighted” images (e.g., T1-(w) weighted or T2w) allow for tissue characterization on the respective MR images by providing different image contrast. Proton density (PD) weighting is based upon the number of H1 nuclei in a respective tissue [3]. Two fundamental parameters—time to echo (TE) and repetition time (TR)—are key to creating image contrast. TR (unit: milliseconds) is the amount of time between successive pulse sequences applied to the same slice. TE (unit: milliseconds) is the time between the application of the RF pulse and the receipt of the peak of the detected echo signal. Both parameters affect the contrast on MRI images: an image with a short TR and TE corresponds to a T1w image (water appears dark and fat bright). A long TR and TE correspond to a T2w image (water and fat appear bright). A long TR and a short TE correspond to a PDw image (with image characteristics of both, T1 and T2) [2]. The sequences used in MRI are summarized in a simplified, more application-related manner in the following section to enhance reader comprehension. Gadolinium contrast media is a paramagnetic metal injected intravenously as a chelated ion in aqueous solution, typically in the form of gadopentetate dimeglumine. At standard extracellular gadolinium concentrations in tissues, T1 effects predominate and therefore gadolinium appears bright on T1 sequences and does not alter the appearance of differently weighted images [4]. For example, gadolinium-enhanced MR sequences are warranted in cases with suspected inflammation or tumors since the detection of inflammation can be improved and tumors can be further characterized. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00002-0 © 2023 Elsevier Inc. All rights reserved.

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B0

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H

H

H

H

H

H

(A)

Net Magnezaon Vector

H

(C)

(B) Mz

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Mxy

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RF on

RF off H20

H20

Fat

Fat

Time

(D)

T1 Recovery

Time

T2* decay

Time

T2 decay

FIGURE 19.1 Basic physics of MRI. (A) As each H1 nucleus rotates around its own axis, its motion induces a magnetic field. (B) Before the H1 nuclei are exposed to an external magnetic field (B0), their axes are randomly aligned. (C) When B0 external magnetization is applied (i.e., the patient is positioned in the core of the permanent active magnet of the scanner), the magnetic axes of the nuclei align with the magnetic axis of B0—some in parallel and others in opposition to it. The cumulative effects of all the magnetic moments of the nuclei sum up to the net magnetization vector. (D) The use of a radiofrequency (RF) pulse creates a flip of the net magnetization vector by a certain angle (α), which leads to two magnetization vector components: transverse (Mxy) and longitudinal (Mz) magnetization. As the transverse magnetization precesses around a receiver coil (there are dedicated surface coils placed on the patient according to the respective anatomical region of interest), a current is induced in that coil. This current leads to the MRsignal. (E) When the RF pulse is turned off, the net magnetization vector of the nuclei realigns with the B0 axis leading to an increase of the longitudinal magnetization (T1 recovery) and a decrease of the transverse magnetization (T2 and T2* decay) [2].

Frequently but not invariably, contrast enhancement of tumors can be considered a sign of malignancy. In these cases, growth of pathological vessels within the tumor are the basis for the contrast enhancement in MRI [5].

19.2

Magnetic resonance imaging sequences

Some basic tenets of MR physics are important for MR image interpretation and are listed in the following. In the vast majority of instances, MR images are acquired with a technique that is referred to as fast spin echo (FSE) image acquisition. With FSE acquisition, fat is bright on all imaging sequences. For this reason, on PD and T2 sequences, water/ edema have bright signal and can be difficult to distinguish from fat (subcutaneous tissue/bone marrow, etc.) [2]. Chemical fat suppression is routinely added to FSE PD and FSE T2 sequences to provide a contrast mechanism in which edema/fluid appear bright and fat-containing tissues (subcutaneous fat, intraarticular fat, marrow fat) are dark. T1 sequence: Water/edema - dark (hypointense) T2 sequence: Water/edema - bright (hyperintense) PD sequence: Water/edema - gray (moderately hyperintense) Fat - bright (hyperintense) on all, T1, T2, and PD Anatomy sequences: T1 and PD Pathology sequences: Fat suppression added to PD or T2 with a long echo time (i.e. fluid-sensitive) and T1 (after intravenous gadolinium injection) 7. Fat suppression is used to suppress signal from fat as a contrast mechanism to profile edema in fatty tissue or allow more facile identification of regions of contrast enhancement. Two techniques are widely used: a. Short tau inversion recovery (STIR) provides intrinsic homogeneous fat suppression even if orthopedic implants are present, but cannot be used postgadolinium. b. Classic chemical fat saturation fat saturation is versatile and can be used postgadolinium but is more prone to artifacts from metal implants and may feature inhomogeneous fat saturation. 8. Intravenous contrast medium (gadolinium) - bright on T1, invisible on other sequences. 1. 2. 3. 4. 5. 6.

In short, there are three different sequences: T1, T2, and PD. Adjusting various settings on the MR scanner for each sequence results in a set of images with signal intensity characteristic for each sequence. The strategy to identify the

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FIGURE 19.2 Examples for main MRI sequences. Coronal pelvis MRI images of a patient without pathology (A C) and with a traumatic injury to the right proximal femur (D and E). The arrows point to the urinary bladder and the arrowheads indicate the edema in the right proximal femur. L, left; R, right.

kind of tissue is based upon the assessment of different signal intensities in the various pulse sequences. This, combined with anatomical knowledge, builds the basis for MRI interpretation. For better understanding, a coronal image of the pelvis is discussed (Fig. 19.2). The content of urinary bladder (arrow) which appears bright (hyperintense) on the T2 sequence (Fig. 19.2C), dark (hypointense) on the T1 sequence (Fig. 19.2A) and gray (moderately hyperintense) on the PD sequence (Fig. 19.2B). This signal behavior is typical for simple fluid, in this case, urine. T1 and PD sequences have better resolution than the T2 sequence and are therefore referred to as anatomy sequences. Consider a traumatic injury to the right hip of the patient (Fig. 19.2D and E), with bone marrow edema due to a contusion. As described earlier, water/edema is bright on T2 fat-suppressed FSE sequences. In fact, fat is very bright in all three sequences, T1, T2, and PD. Due to the bright T2 fat signal in the bone marrow, the bone edema caused by the contusion is also bright on T2 and therefore likely to be masked. Hence, it is reasonable to suppress the fat signal of the bone marrow for the T2 sequence so that bone edema becomes apparent (Fig. 19.2D). The same advantage of fat suppression applies to MRI sequences where a contrast medium (gadolinium) is intravenously injected into a patient. Gadolinium provides bright signal only on T1 sequences and is invisible on T2 and PD sequences. If pathology (e.g., bone tumor or a joint inflammation) takes up gadolinium, the bright signal caused by it might be masked by the bright signal of fat on T1. Therefore we use fat suppression for gadolinium-enhanced T1 sequences (Fig. 19.2E) and for T2/PD sequences to create a contrast mechanism to detect edema in fatty tissues. There are some potential disadvantages of fat suppression in that it reduces the overall resolution of the sequence and metal-related artifacts from implants cause heterogeneous fat suppression (among other things), that makes interpretation of signal challenging.

19.3

Magnetic resonance imaging versus computed tomography

Both computed tomography (CT) and MRI show cross-sectional images of the body. The underlying technologies of these modalities, however, are entirely different. A CT scanner basically consists of an X-ray tube that rotates 360-degrees around the patient acquiring images of the body while the patient is positioned on a moveable table. The beam passes through the respective body region and is attenuated to a different extent by the density of the tissues therein. The beam then strikes the sensor system and is eventually used to produce 2D and 3D datasets. In contrast to the attenuation principle of a CT scanner, for MRI, the patient is positioned in a strong magnetic field as mentioned earlier. While in that field, radiofrequency pulses are applied for a short time to excite hydrogen atoms in the body. By monitoring the excitation and subsequent relaxation of the hydrogen atoms, signal is generated and registered. After several complex technical steps (see the earlier sections of this chapter), 2D and 3D datasets are created [2].

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The main differences between both techniques for routine musculoskeletal imaging are the following: 1. CT uses ionizing radiation and contributes 10% 15% to the overall annual dose per capita [6]. 2. CT is excellent for the depiction of bone but lacks detailed soft tissue contrast which is a strength of MRI. Further, new MRI techniques for radiation-free bone imaging are promising [e.g., ultrashort TE (UTE) MRI sequences [7]] Please see section Areas of future research at the end of the chapter for more details. 3. While CT scans last only a few seconds and are generally well tolerated, MRI can last up to 30 40 minutes, and patients are required to keep very still. However, new MRI applications (e.g., parallel imaging/compressed sensing) are allowing increasingly shorter scan times (see Section 19.10). 4. Back pain due to patient position during the MRI scan and claustrophobia due to the narrow gantry makes a significant number of scans impossible (approxiamately 10% of all cases). 5. A long list of contraindications for MRI (e.g., dated pacemaker, metal implants or splinters, pregnancy) has to be screened for by the referring physician prior to scheduling. There are pacemaker databases online available summarizing MRI compatible and noncompatible pacemaker devices [8].

19.4 Magnetic resonance imaging appearance of musculoskeletal tissue—normal and pathology 19.4.1 Short T2 tissues The tissues of the human body can be divided into those that are “visible” on standard clinical sequences and those that are “invisible” because their mean echo time (TE) is too short to provide a detectable signal. All tissues contain a mixture of short and long T2 components in a certain proportion. The invisible tissues have a majority of short T2 components and a minority of long T2 components. The short T2 components typically do not provide enough signal to be detectable in relation to image noise levels. Healthy tendons, ligaments, and cortical bone are comprised of tissue with low mobile proton content (scarce water content) resulting in dark signal on most sequences (T1, T2, and PD). Tendons or ligaments affected by degeneration feature an increase of long T2 components. Degeneration leads to disorganization of collagen fibers and higher water content, presenting with a brighter signal on all standard MRI sequences.

19.4.2 Tendons There is a lack of consistent terminology for the different forms of pathologic tendon conditions. Tendinopathy (syn. tendonopathy) is defined by many as an umbrella term for all kinds of tendon pathologies [9]. Tendinosis: Usually implies intratendinous degeneration of a tendon’s collagen in response to chronic overuse without significant inflammation [10]. Tendinosis is often accompanied by thickening or thinning and rounding of the tendon. The MRI signal of the affected tendon can become intermediate bright (gray) on the T1, T2, and PD sequence, but not as bright in T2 as water (in distinction to tendon tears) (Fig. 19.3). Tears: Tears can be categorized as a complete or partial tear with either a predominately horizontal or longitudinal orientation. Tears have a bright signal on T2 images. Unfortunately, these features are indistinguishable from “mucoid degeneration.” Histologically, mucoid degeneration features mucoid patches and vacuoles between thinned collagen fibers causing the affected region to soften with alteration of the biomechanical properties [11]. Tears can be divided into partial tears (i.e., partial discontinuity) and full thickness tears (i.e., complete transection). Tenosynovitis: If signs of tendinitis are accompanied by bright fluid in the tendon sheath (if present) visible in T2 and PD sequences, tenosynovitis should be suspected. Just as the intestinal mesentery is formed, the layers of the tendon sheath form a thin low-intensity band “mesotenon” which becomes apparent when effusion is present. A minimal amount of fluid in joints and tendon sheaths can be physiological. A pathologic effusion in these compartments is not diagnosed by a different MRI signal characteristic of the fluid but by its increased amount. If the tendon does not communicate with a joint, then no fluid should be present in the tendon sheath in a healthy state. In contrast, the sheath of the flexor hallucis longus tendon sheath communicates with the tibiotalar joint and thereby might be filled with some joint fluid physiologically. After intravenous contrast injection, inflamed tissue shows avid contrast uptake on T1 fatsuppressed sequences [12]. Other nonpathologic causes of increased signal within tendons are tendons with an intratendinous fibrocartilaginous sesamoid (e.g., typically in the posterior tibial tendon) [13], and where the tendon normally fans out or merges with other tendons. The increased signal in these cases is different enough from the appearance of fluid in tear and tenosynovitis to generally allow for distinction.

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FIGURE 19.3 Tendinosis of posterior tibial tendon. The posterior tibial tendon shows intermediate signal below the medial malleolus on both PD [(A) coronal plane and (C) sagittal plane] and fat-saturated T2 [(B) coronal plane and (D) sagittal plane] sequences. In distinction to tendon tears, tendinosis features no T2 signal that is as bright as water.

19.4.3 Pseudo tendon pathology—the magic angle effect or phenomenon The magic angle effect is an important phenomenon in musculoskeletal MRI, which if unaccounted for, can erroneously suggest tendon pathology. Inherently, this artifact appears primarily on sequences with a short TE (less than 32ms; e.g., T1, PD, and gradient echo sequences) and is confined to regions of tightly bound collagen which are oriented at 54.74 from the main magnetic field (i.e., the long axis of the patient table of the MRI scanner). The phenomenon produces an artificially bright signal, thus potentially being mistaken for tendinosis [14] (Fig. 19.4). Typically affected sites include the peroneal tendons as they traverse the lateral malleolus and the cubital groove. Comparison with longer TE sequences (e.g., T2 including FSE T2) can be helpful since these sequences are less susceptible to the magic angle effect. Furthermore, in cases with questionable tendon pathology around the ankle, patients can be positioned in the prone position with plantarflexion to eliminate the potential magic angle artifact by changing the critical angle of the tendons to the main magnetic field.

19.4.4 Ligaments Ligaments have a similar MRI appearance as tendons. However, some ligaments frequently feature fatty streaks, which have a bright signal on nonfat-saturated T1, T2, and PD sequences. This appearance should not be misinterpreted as partial tears. Scarring of acutely injured ligaments starts after a few days and may lead to underestimation of the injury in MRI performed in the subacute setting. Scarred ligaments are commonly thickened and have lost their regular multifascicular appearance (Fig. 19.5). Mass-like thickening after ligament injury can result in impingement syndromes [15].

19.4.5 Bone Cortical bone, almost lacking in mobile protons, appears black on T1, T2, and PD sequences. In comparison, cancellous bone is predominately filled with yellow fat in adults, which is bright on T1, T2, and PD sequences. After the age of 25 years, red bone marrow with active hematopoiesis is mainly located in the axial skeleton but can also be found scattered in the proximal long bones. Red bone marrow features a T1 signal less intense than normal muscle and yellow bone marrow due to the reduced fat content and appears brighter on fat-saturated T2 sequences.

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FIGURE 19.4 The magic angle artifact of the peroneal tendons. (A) Sagittal (PDTR 3500, TE 30), magic angle artifact of the peroneal tendons is typically seen around and below the lateral malleolus (green arrowhead), suggesting tendinosis by increased signal of the tendon at an angle of 54.74 from the main magnetic field (B0) on sequences with a short TE. (B) Sagittal PD with 20-degree dorsiflexion; reduction of magic angle artifact of the peroneal tendons. (C) Sagittal PD with the prone position; no magic angle effect of the peroneal tendons is noted.

FIGURE 19.5 MRI appearance of scarred versus normal ligaments. Scarring of the deltoid ligament due to chronic deltoid ligament injury is indicated by the red arrows (A, PD) and (B, fat saturated T2). In comparison (C), the intact deltoid ligament features a normal fan-shaped appearance owing to internal fat striations.

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Bone contusions create an even brighter signal on fat-saturated T2 sequences than red bone marrow and are comprised histologically of a mixture of edema, hemorrhage, and necrosis. Inherently, avulsion fractures do not create marked bone marrow edema. Even nondisplaced fractures are frequently surrounded by extensive edema, but the fracture line itself is best visible as a black line on T1 images that traverses the trabecular bone. Insufficiency and fatigue fractures represent other mechanisms of bone injury. Insufficiency fractures are the result of normal stresses applied to abnormal bone (quantity or quality of bone) and fatigue fractures, which are due to abnormal stresses applied to normal bone (e.g., march fractures). Pathological fractures are generally the result of focally diseased bone (e.g., tumors, both malignant and benign). In stress fractures, the occurrence of a particular fracture line is preceded by an edematous stress reaction that is exclusively visible on MRI as T2 and PD hyperintense signal in the involved bone marrow as well as the adjacent soft tissues [16]. Conservative treatment including reduction of mechanical stress at this stage can prevent progression to a complete fracture. In patients with rheumatologic workup, the unique ability of MRI to detect bone marrow edema is crucial since bone marrow edema typically precedes erosions that are not visible on X-rays until several months or years have passed and, thereby, allows much earlier diagnosis and treatment initiation before irreversible damage occurs.

19.4.6 Hyaline and fibrocartilaginous cartilage Hyaline cartilage, covering the articulating surfaces of most joints, is frequently assessed by PD and T2 sequences due to its high water and collagen content featuring depth-dependent signal characteristics [17]. Superficial cartilage layers have a brighter signal on T2/PD sequences with lower signal closer to the calcified cartilage layer. This zonal variation in signal intensity has been referred to as grayscale stratification. There are also other types of sequences available for imaging of hyaline cartilage [18]. However, fibrocartilaginous cartilage, e.g., found in the meniscus and the articular disk of the temporomandibular joint, are short T2 tissues as previously noted and have low signal on all standard imaging sequences. In general, the signal of hyaline and fibrocartilaginous cartilage increases on T2 and PD sequences when collagen fiber derangement and swelling of the cartilage occurs. The signal alteration of cartilage with or without cartilage substance loss is referred to as chondromalacia. Owing to repeated microtrauma, osteochondritis dissecans, i.e., a separation of a chondral or osteochondral fragment with the gradual fragmentation of the articular surface, can occur. Frequently, this leads to intraarticular loose bodies that follow the same MRI signal characteristics as the original tissue.

19.4.7 Muscle Healthy muscle tissue has a T1 and T2 that is characterized with a signal brightness that is intermediate in comparison to fat and cortical bone. The MRI appearance of normal muscle varies with respect to the degree of internal fat, relationship to the tendon, and type of attachment. In acute muscle injury grade I, there is focal or diffuse high signal intensity with feathery appearance to the muscle on fluid-sensitive sequences images without disruption of muscle fibers. In grade II, the musculotendinous junction is partially torn, and areas of inhomogeneous hematoma can frequently be seen. The musculotendinous junction is the site of connection between tendon and muscle. In this region, the force generated by muscle contraction is transmitted to the tendon, making this weakest element of the muscle-tendon complex susceptible for injury. Partial tears are frequently accompanied by a space-occupying hematoma that can have various appearances on MRI sequences depending on the stage of blood degradation. In grade 3, complete disruption of the musculotendinous junction occurs with extensive edema and hemorrhage and wavy tendon morphology [19]. Edema of the entire muscle belly can be an early sign of denervation (e.g., nerve injury or nerve entrapment) or nonusage of the muscle belly (e.g., in the case of a full tendon tear the remaining muscle belly has no counterpart to contract against). After a few weeks to months of insufficient mechanical use or denervation, the muscle starts to atrophy. T1-bright fatty streaks within the muscle are signs of fatty atrophy and are considered irreversible changes (Fig. 19.6).

19.4.8 Bursa/synovia Healthy bursae are sliding, sparsely fluid-filled layers that are frequently not visible on MRI. Only minimal T2 bright fluid can be considered normal. A greater amount of fluid, irregular contour, or septations within the bursa are signs of

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FIGURE 19.6 Denervation edema. Denervation edema of the abductor hallucis brevis (yellow dashed outline) and flexor digitorum brevis muscles (orange dashed outline) is caused by compression of the medial branch of posterior tibial nerve (arrow) by a talocalcaneal coalition with bony outgrowth of the sustentaculum tali and the corresponding talus. (A) Coronal plane through midfoot, (B) coronal plane through hindfoot, (C) transverse plane.

inflammation. Bursal inflammation is a nonspecific finding and can be seen with mechanical irritation, through rheumatologic disorders, or infection. Similarly, healthy synovia, which is the inner lining of synovial joints, is very challenging to visualize on MRI. When mechanical irritation of a joint occurs, e.g., activated osteoarthritis, or joints are affected by inflammation, e.g., in rheumatologic conditions, joint fluid is frequently present. In the acute setting, the synovial hypertrophy can be difficult to distinguish from the appearance of simple fluid on the PD or T2 sequence. Both conditions, inflamed bursa and synovial joints, show extensive gadolinium uptake on T1 sequences.

19.5 Tailored magnetic resonance imaging protocol for the foot and ankle— indication driven 19.5.1 Imaging orientation Defining the scanning protocol of the ankle and foot can be challenging. Preferably, the foot should be placed in 20 degrees of dorsiflexion during scanning to decrease the magic angle phenomenon noted previously [20] while retaining the anatomical presentation of the foot. Of note, diagnostic capabilities are markedly enhanced when ankle and foot MRI are considered two distinct studies.

19.5.2 Tailored magnetic resonance imaging protocols There are routine MRI protocols with imaging parameters for the ankle and foot (Table 19.1). Frequently, each institution defines a basic scan protocol for the particular anatomical region (most common: ankle or foot or foot & ankle) depending on the clinical question. This choice of the basic protocol takes into account that all planes should be covered and that at least one sequence should be sensitive to altered signal in fatty tissues [generally intrinsically (inversion recovery) or chemically fat-saturated (PDw or T2w)] to increase the visibility of bone marrow or soft tissue edema and geodes, for example.

TABLE 19.1 Ankle and foot magnetic resonance imaging (MRI) scanning protocol. MRI parameter Plane and sequences

TR/TE (Ms)

ST/space (mm)

FOV (cm)

Matrix

ETL (count)

Acquisition time (min) d

5000/33

3.3/0.5

11 15

512 3 256

12

2:25

Coronal FSE PDw

5000/33

4.0/0.5

11 15

512 3 384

12

3:28

a

4500/33

3.0/0.5

11 15

512 3 384

12

3:05

4000/13

3.3/0.5

18

256 3 192

12

6:49

3500/50

3.3/0.5

11 15

512 3 256

12

5:22

Coronal FSE T2w FS

5000/45

4.0/0.5

11 15

512 3 384

12

5:25

c

3500/50

3.0/0.5

11 15

512 3 384

12

7:42

2675/35

2.5/0.4

11 13

512 3 512

8

4:28

Coronal FSE PDw

3200/30

2.2/0.2

15 17

512 3 512

8

3:06

a

3450/30

2.5/0.3

15 17

512 3 512

8

2:53

5000/77

2.5/0.3

11 13

512 3 512

14

3.29

Coronal FSE T2w FS

3475/60

2.2/0.2

11 13

512 3 512

16

4:53

b

3100/60

3.0/0.5

11 15

512 3 512

14

2:42

Ankle Without fat saturation Axial FSE PDwa a

Sagittal FSE PDw

With fat saturation Sagittal FSE-IR (IR170)a Axial FSE T2w FSb b

Sagittal FSE T2w FS Foot

Without fat saturation Axial FSE PDwa a

Sagittal FSE PDw

With fat saturation Axial FSE T2w FSb b

Sagittal FSE T2w FS

ETL, Echo train length; FOV, field of view; FSE, fast spin echo; IR, inversion recovery; PDw/T2w, “w” is the abbreviation for “weighted.”—weighted; ST, slice thickness; TE, time to echo; TR, repetitive time. a Basic protocol should include these four sequences for the ankle and these three sequences for the foot. b Possible additional fat-suppressed fluid-sensitive sequences for evaluating ankle and foot structures. However, fat suppression has a longer scan time as a trade-off. c Possible additional sequences that can be used to replace Sagittal FSE-IR. Sagittal FSE T2w FS sequence has a higher resolution but is more prone to magnetic field inhomogeneities and susceptibility artifacts. d Scan time was calculated on 3.0 T GE Diskovery 750; GE Healthcare, Milwaukee, WI.

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19.5.3 Optimized imaging planes For specific clinical indications, dedicated imaging planes have been introduced in the literature. Peroneal view: An oblique sagittal plane that runs parallel to a line between the lateral border of the posterior portion of the calcaneus and lateral margin of the peroneal tubercle at the mid-calcaneal level has been termed the “peroneal view.” The sequence features a 2 mm slice thickness for improved evaluation of peroneal tendon pathology [21]. This results in an image that profiles the peroneal tendons along their long axis, through their 90-degree directional change (Fig. 19.7). Syndesmotic view: An oblique-axial plane, “syndesmotic view” is acquired at approximately 45 degrees (tilted toward anterior) orthogonal to the course of the ligaments and offers an improved interpretation of ligament continuity, thickening, and contour [22] (Fig. 19.8).

FIGURE 19.7 Peroneal view. The true sagittal PD fat-saturated image (A) shows the peroneus brevis tendon, which is not continuously visible as the tendon has a slightly oblique course. In contrast, the sagittal-oblique sequence in (B) [see reference lines in (C)] shows the entire course of the tendon to the attachment at the base of the fifth metatarsal and depicts bright signal alterations (arrows) of the tendon due to a partial tendon tear which was confirmed on axial-oblique PD sequences (not shown).

FIGURE 19.8 Syndesmotic view. The axial PD nonfat-saturated image (A) shows that the anterior syndesmotic ligaments appear discontinuous due to the oblique course of the ligament. On the axial oblique PD sequence (B), its integrity is clear (arrow). The reference lines are visible in (C).

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Calcaneofibular ligament view: For improved evaluation of the calcaneofibular ligament of the lateral collateral band complex, a coronal oblique view with a 55-degree angle toward the plantar fascia can be performed [23]. In cases with a suspected Morton’s neuroma (see Section 19.9), it is essential to examine the patient in the prone position with the ankle in extension since the pressure created in the interdigital space improves detection of the neuroma significantly [24].

19.5.4 Metal artifact reduction sequences Imaging around metal and postoperative imaging present substantial challenges for MRI of the ankle. The artifacts encountered are caused by differences in magnetization between human tissue and metal prostheses. Four general types of artifact include: G G G G

Loss of signal intensity Signal-intensity pileup Geometric distortion Failure of fat suppression.

Basic MRI principles and consideration of magnetism, dephasing, and signal misregistration, and the properties of the metal in question promote parameter changes in conventional MRI sequences that are broadly referred to as metal artifact reduction sequences. Because magnetization and inhomogeneity are directly proportional to field strength (magnetic field strength is given in units of Tesla (T) e.g., 3 T), a greater artifact is expected in a 3 T magnetic field than in a 1.5 T magnetic field. In general, a higher field strength allows higher resolution because more signal can be received. Concomitantly, susceptibility artifacts from metal implants are generally considered to increase linearly with field strength. Furthermore, several implanted older medical devices like cardiac pacemakers and stapes prostheses are only MRI compatible up to 1.5 T. There are several publications summarizing the most suitable metal artifact reduction sequences [25,26]. The area in close proximity to the metallic hardware is of main interest since stress reactions, fractures, as well as bone resorption and aseptic loosening, occur here. However, that is precisely the area that is most frequently affected by the artifacts. To tackle these problems several dedicated sequences have been developed. The multiacquisition variable-resonance imaging combination (MAVRIC) is a multispectral sequence based on 3D FSE acquisition by using a series of frequency-selective excitations with variable radiofrequency impulses to reduce in-plane and through-plane artifacts. Sequences with slice encoding for metal artifact correction (SEMAC) reduces in-plane and through-plane artifacts by using additional z-directional phase-encoding steps. Both MAVRIC and SEMAC sequences have similar capabilites for metal artifact reduction (Fig. 19.9). Currently, there is a hybrid sequence that combines multispectral and z-directional phase-encoding step, namely, MAVRIC-SL [25].

19.6

Magnetic resonance imaging anatomy of the foot and ankle

19.6.1 Ankle On the lateral site, the ankle is stabilized by the syndesmosis and the lateral ligament complex, which is located just inferior to the syndesmosis.

19.6.1.1 Ligaments The anterior inferior tibiofibular ligament (AITFL) of the syndesmosis runs from the anterior/superior lateral tibia posterior/inferior to the fibula (B 40 degrees). Therefore it is commonly not visible in continuity on one axial slice. Oblique-axial “syndesmotic view” sequences provide a solution. Bright fatty streaks on T1- and T2 nonfat-saturated sequences within the healthy AITFL are commonly seen. In ankle injuries, the AITFL has been seen to tear (1% 11% of cases) whereas the posterior inferior tibiofibular ligament (PITFL) remains intact [27]. The third primary component of the syndesmosis, the interosseous membrane, strongly connects the tibia and the fibula and divides the anterior from the deep posterior compartment of the lower leg. In more severe injuries, an avulsion fracture of the posterolateral tibia at the attachment of the PITFL occurs, called “Volkmann’s triangle.” The anterior talofibular ligament (ATFL) and posterior talofibular ligament (PTFL) of the lateral ligament complex are visible on axial images. Similar to the AITFL, the PTFL commonly contains fatty streaks when healthy. The calcaneofibular ligament has an oblique-axial course beneath the peroneal tendons and, as a unique feature, spans two joints, both the high

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FIGURE 19.9 Example demonstrating MRI sequences original and with metal artifact reduction. Hip replacement prosthesis causes marked artifacts in proximity to the prosthesis on axial (A) and coronal (C) T1w sequences hampering the assessment of the structures nearby. In contrast, on axial SEMAC (B) and coronal MAVRIC (D) sequences the metal artifacts are significantly reduced.

FIGURE 19.10 Tear of calcaneofibular ligament. This example shows a tear of the calcaneofibular ligament with consecutive dislocation of the peroneal tendons into the gap (A, T2w sequence) preventing healing of the ligament tear. The STIR image (B) shows edematous soft tissue structures adjacent to the tear.

ankle and subtalar joint. Injuries to the lateral ligament complex always follow the sequence that the ATFL tears first, then the calcaneofibular ligament if higher forces are involved. Only very rarely and in severe cases will the PTFL tear [28]. Notably, the rupture of the calcaneofibular ligament can cause dislocation of the peroneal tendons into the gap between the ligament ends and, thereby, lead to delayed or incomplete healing of the calcaneofibular ligament (Fig. 19.10).

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FIGURE 19.11 Osteochondritis dissecans of the talar dome. On the coronal PD nonfat-saturated image (A) the lateral talar dome shows an osteochondral fragment (long arrow) after an acute mechanical trauma three months earlier. Cartilage, as well as bone signal, is seen within the fragment, and the fat-saturated fluid-sensitive T2wFS sequence (B) shows edema and cystic transformation (short arrow) within the residual talar bone. The fragment is surrounded all around by PD- and T2-bright joint fluid indicating an intraarticular loose body.

The medial collateral (deltoid) ligament of the ankle is comprised of deep and superficial layers. Ankle trauma is quite common and frequently merits MRI evaluation.

19.6.1.2 Bony defects and ligament injuries Acute chondral and osteochondral defects are more commonly seen at the lateral talar dome (Fig. 19.11), whereas osteochondritis dissecans, caused by chronic (i.e., degenerative) microtrauma, are instead located medially [29]. In ankle injuries with suspected significant ligament injuries, MRI should be performed as soon as possible since accurate detection of the extent of injuries can be underestimated due to scar tissue formation that bridges the ends of ligament tears within 10 20 days after trauma already [30].

19.6.2 Hindfoot 19.6.2.1 Tendons Primarily, four tendon groups are responsible for the movement of the foot and the majority of pathologies of these tendons manifest at the hindfoot. Dorsiflexion of the ankle is generated via the tibialis anterior, extensor hallucis longus, and extensor digitorum tendons which are rarely affected by pathologies [31]. Of these, the tibialis anterior tendon is occasionally affected by chronic tendinosis at the distal insertion at the medial cuneiform and base of the first metatarsal bones (appearing thickened with intermediate signal on T1, T2, and PD sequences) and tearing occurs between the upper and lower retinacula on the level of the talocrural joint line (tears appear bright as water on T2 in distinction to tendinosis). Plantarflexion of the foot is generated through the actions of the tibialis posterior, flexor digitorum longus, flexor hallucis longus, and Achilles tendons. The tibialis posterior tendon cross-section should be oval and twice as thick as the adjacent flexor digitorum longus tendon. Rupture of the tibialis posterior tendon occurs commonly posteriorly to the medial malleolus owing to a localized hypovascularity and the significant friction possible at this site. As one of the main dynamic stabilizers of the longitudinal arch of the foot, chronic tendinosis of the tibialis posterior tendon

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eventually may contribute to pes planus. Typically, MRI reveals reactive marrow edema under the tibialis posterior tendon groove caused by posterior tibialis tendon dysfunction (PTTD) [32]. Since the flexor hallucis longus tendon communicates with the ankle joint, fluid in the tendon sheath is normal and not a sign of tenosynovitis, as would be the case for other tendons. In the healthy population, the flexor hallucis longus tendon is rarely injured or inflamed. In comparison, flexor hallucis longus tendon pathology is much more common in dancers, leading to the moniker “Achilles tendon of the dancer” [33]. The normal Achilles tendon has a sagittal width of less than 1 cm and has a concave or flat shape of the ventral surface on axial images. Achilles tendon tendinosis leads to T1, and T2 increased signal with rounding and thickening of the tendon predominately in the middle third of the Achilles tendon (Fig. 19.12). Achilles tendon rupture almost exclusively occurs in tendons with underlying chronic structural damage approximately 6 cm above the calcaneal insertion at the watershed area [34]. Eversion of the foot is generated by the peroneal tendons. An avulsion fracture at the site of the peroneal brevis tendon at its distal insertion at the base of the fifth metatarsal can be easily overlooked on MRI. As previously noted, avulsion fractures are often not accompanied by bone marrow edema and small bony flakes, and are therefore

FIGURE 19.12 Achilles tendon tendinosis. The Achilles tendon of this 65 years old overweight woman shows increased signal on the sagittal T1w (A), sagittal STIR (B), and axial fat-saturated T2 images (D) as well as rounding and thickening of the tendon with a convex anterior boundary (C).

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FIGURE 19.13 Split tear of peroneus brevis tendon. Due to partial volume artifact, the split tear of the peroneus brevis tendon is not visible on the axial PDw image (A) but on the coronal oblique PDw image (arrow in B). Please see (C) for reference lines.

inconspicuous. If the peroneal tendons rupture, in particular, the peroneus brevis tendon, the tear is commonly partial and vertical in shape (“split tear”). The split tear leads to a pathologic C-shape on axial T2 and PD images (Fig. 19.13). However, there is a normal variant with a doubled peroneus brevis tendon that can be identified by scrutinizing the corresponding muscle bellies. In the case of a common variant, each tendon has a separate muscle belly. Postgadolinium T1 fat-saturated sequences are helpful in diagnostically these challenging cases as split tears, contrary to the normal variant, show contrast enhancement.

19.6.2.2 Enthesitis—plantar fasciitis Plantar fasciitis is the most common cause of inferior heel pain and originates most likely from altered biomechanics (e.g., flattened longitudinal arch or increased Achilles tendon tension), leading to overuse and microtrauma with subsequent inflammation at its insertion [35]. Plantar fasciitis is a clinical umbrella term that may relate to several structural changes at the plantar fascia proximal attachment: MRI signs of plantar fasciitis are a T2-very bright tear or plantar fascia thickness greater than 5 mm, altered signal intensity within or adjacent to the plantar fascia, and a bony spur or bone edema at or above the plantar fascia insertion [36] (Fig. 19.14). Thickening of the plantar fascia is frequently seen in plantar fasciitis but is not a mandatory condition to make the diagnosis. Note, that small bony spurs can be invisible on conventional MRI sequences due to their low mobile PD content and when in doubt, a lateral ankle X-ray should be performed. A recent study showed that the size and morphology of the tuber calcanei might influence the occurrence of plantar fasciitis [37]. Furthermore, the plantar fascia is frequently affected in rheumatologic patients by an extra-axial manifestation of seronegative spondyloarthropathies showing inflammation [35]. MRI, contrary to all other imaging modalities, features the unique ability to detect bone marrow edema using fat-saturated fluid-sensitive T2 or PD sequences. Bone marrow edema is known to cause pain and to precede bony erosions that are visible on x-rays by several month or years [38].

19.6.3 Midfoot 19.6.3.1 Degenerative joint disease The midfoot comprises tarsal bones which, especially in the elderly and overweight, frequently shows classic radiological signs of osteoarthritis that can be assessed on MRI even more accurately than on radiographs. Most commonly the talonavicular and the calcaneocuboid joint are affected by degeneration. One hallmark of osteoarthritis is joint space reduction, assessed on radiographs as an indirect measure of the articular cartilage height, and this can be assessed directly on T2 and PD MRI sequences. Furthermore, cartilage delamination, fissures, and cartilage swelling, often associated with pain, cannot be detected on radiographs since they do not affect the width of the joint space. Osteophytes frequently contain T1 bright fatty bone marrow surrounded by a more

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FIGURE 19.14 Plantar fasciitis. The plantar fascia is comprised of a smaller lateral (arrow) and a stronger central cord. In this case, the central cord at the calcaneal insertion is markedly enlarged with a high signal (arrowheads) on sagittal fat-saturated T2 (A) and coronal PD (B) images.

FIGURE 19.15 Navicular stress fracture. This example shows a stress fracture of the medial aspect of the navicular bone with a thin fracture visible as a dark hairline on the T1w image (A) and adjacent bone marrow edema on the fat-saturated T2 image (B). This case shows an unusual rather horizontal orientation of the fracture line. In navicular stress fractures, the fracture line is frequently orientated in the sagittal plane.

or less thin cortical bone featuring dark signal on all conventional MRI sequences. Subchondral sclerosis can be seen as T1 dark thickened subcortical bone on all conventional sequences. Subchondral cysts (i.e., geodes, well-defined lytic lesion in the periarticular surfaces) are visible as bubbly T2 and PD strongly bright subchondral structures occasionally surrounded by fluffy T2 and PD bright edematous bone marrow changes. Geodes can become relatively large and frequently show full or rim-enhancement on T1 postgadolinium sequences, and is therefore occasionally misdiagnosed as a tumor sinister [39].

19.6.3.2 Stress fractures Particularly in runners, fatigue fractures of the navicular bone as the integral structural link for force transmission between the midfoot and the hindfoot occur frequently [40]. Navicular stress fractures are typically oriented in the sagittal plane and are therefore best visible on axial and coronal images. In the beginning, stress fractures manifest only as areas of bright marrow edema on fat-saturated T2 and PD sequences, and at later stages, may exhibit a line of lowsignal intensity best visible on T1 sequence representing a fracture line that can be partial or complete (Fig. 19.15).

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Complete fractures bear the risk of pseudoarthrosis and osteonecrosis of the lateral navicular segment which is relatively poorly vascularized.

19.6.4 Forefoot 19.6.4.1 Morton’s neuroma Patients with a Morton’s neuroma typically present with forefoot pain which worsens with activity. The neuroma is comprised of perineural fibrosis around the plantar interdigital nerve and most frequently located in the second or third interdigital space. Examination of these patients in the prone position with plantarflexion improves detection of Morton’s neuroma [24]. Since perineural fibrosis can be found to some extent also in the asymptomatic population, a T1 rather dark and T2 rather dark or moderately bright tissue mass with a mediolateral dimension of more than 5 mm located plantar to the interdigital space is considered as a Morton’s neuroma (Fig. 19.16) [41].

FIGURE 19.16 Morton’s neuroma. The location of the intermetatarsal neurovascular bundle is visible (A). In the third interdigital space a T1w hypointense (B and C) and T2w hyperintense (D) dumbbell/ovoid-shaped lesion is visible.

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FIGURE 19.17 Osteomyelitis. There is a mild hypointense signal (arrow) in the calcaneus on the T1w image (A) with a corresponding bright signal (arrow) on the fat saturated T2 image (B) indicating replacement of the natural fatty yellow marrow by osteomyelitis. The same area shows vivid contrast enhancement (arrow) on the fat-saturated post-gadolinium image (C). A concomitant soft tissue infection (arrowheads) due to an ulcer (not shown) caused by diabetic neuropathy is frequently seen.

19.6.4.2 Osteomyelitis Osteomyelitis is infection within the bone. It occurs frequently in diabetic patients that suffer from peripheral neuropathy and peripheral vascular disease. In the immunocompetent patient, the vast majority of osteomyelitis occurs from skin ulceration [42]. MRI findings of osteomyelitis include a low signal on T1 and increased signal on fluid-sensitive sequences in the bone adjacent to the soft tissue ulceration (Fig. 19.17).

19.6.4.3 Plantar plate tears The plantar plate is a plantar-sided reinforcement of the joint capsule connecting the base of the metatarsal heads to the base of the proximal phalanx and, therefore, an important stabilizer of the metatarsophalangeal joint. On the plantar side, it is connected with the flexor tendon sheath. It is hypointense (dark signal) on all conventional MRI sequences. During extreme traumatic dorsiflexion injury of the toe may result in a sprain or detachment of the plantar plate (i.e., "turf toe") [43]. This appears as a bright signal in and around the plantar plate on fat-saturated PD and T2 sequences with preserved black fibers of the plantar plate, or a partial or complete tear visible as signal on these sequences as bright as water (Fig. 19.18). Rupture occurs typically near the insertion to the base of the proximal phalanx, and the tear features a transverse course. Plantar plate injuries are typically located at the second ray and injuries to the first ray are called “turf toe” injuries and is commonly seen in sports on artificial surfaces [43].

19.7

Areas of future research

Improvements in modern imaging have led to progress in understanding and treatment of pathologies of the complicated system that is the foot and ankle. There are several promising MR imaging techniques emerging that will be addressed briefly:

19.7.1 Radiation-free bone imaging Today, the use of ionizing radiation by X-rays and CT is required to depict anatomy or pathology of bones because conventional MRI cannot show these structures sufficiently. The reason for this limitation is that bony structures have very few hydrogen molecules that are excitable electromagnetically. Therefore sufficient signal cannot be acquired from bones and they remain black (hypointense) like other similar tissues (e.g., tendons and ligaments). Hence, these tissues cannot be distinguished by MRI since they lack in contrast between each other. However, new MRI techniques summarized as UTE MRI sequences can acquire the signal from these tissues much faster (ultrashort  , 0.0001 seconds) allowing detection from even these tissues lacking in excitable molecules (Fig. 19.19) [7]. On MRI, tissues are composed of “short T2” and “long T2” components. Bone, meniscus and scar tissue, for example, are comprised of mainly “short T2” components with a signal intensity that drops much faster after the

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FIGURE 19.18 Plantar plate tear of digitus I, “turf toe.” The first metatarsophalangeal joint is shown. The intact plantar plate is a fibrocartilaginous anatomical structure (arrowheads in A) with proximal insertion at the metatarsal head and distal insertion at the base of the proximal phalanx. It overlies the flexor hallucis longus tendon (arrows in A and B). Rupture of the plantar plate at the first metatarsophalangeal joint (C) which is referred to as “turf toe.”

electromagnetic excitation (orange dashed line) than tissues like muscle or fat comprised of mainly “long T2” components (yellow line). Conventional MRI sequences use an echo time (TE) . 10 Ms whereas UTE sequences use a TE , 1 ms. Using conventional MRI sequences, the signal intensity of these “short T2” components has almost completely decayed whereas there is sufficient signal detectable using UTE sequences.

19.7.2 Magnetic resonance imaging scan time reduction Reduction of scanning time is desired since patients do not tolerate being in the gantry of the scanner for more than 30 40 minutes very well, especially in supine position with limited ability to move. Furthermore, scanning time is very expensive and in most healthcare systems there are fixed rates for an MRI of a respective body region of interest no matter the duration of the examination. Therefore the goal of the healthcare provider is to shorten the MR protocols as much as it is reasonably achievable. Reduction of scan time can be achieved by several measures of which parallel imaging and compressed sensing are widely used. Parallel imaging is a widely used technique where the known placement and sensitivities of receiver coils are used to assist spatial localization of the MR signal. Having this additional information about the coils allows reduction in number of phase-encoding steps during image acquisition. This, in turn, potentially results in a several-fold reduction in imaging time.

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FIGURE 19.19 Physical background of conventional versus ultrashort time to echo (UTE) MRI sequences.

19.7.3 Volumetric, isotropic 3D magnetic resonance imaging sequences Volumetric, isotropic 3D FSE sequences with submillimeter slice thickness have been developed, and these offer interesting possibilities for workflow efficiency, as well as the ability to reconstruct the images in any plane, at variable slice thicknesses. For example, ligaments of the ankle are frequently oriented obliquely to all three orthogonal planes, and reconstruction of planes paralleling the course of the ligaments makes the assessment of those structures more accurate. Names of these 3D sequences vary by vendor, e.g., SPACE, CUBE, and VISTA [44]. In the light of the recent improvements of MRI technology, there is a tendency to perform one 3D sequence that can replace acquisition of the three orthogonal planes. In summary, the longer scanning time needed for acquisition of a high-resolution isotropic 3D sequence is outweighed by the acquisition of each of the three planes separately, which would require even more time.

References [1] Bloch F. The principle of nuclear induction. Science 1953;118:425 30. [2] Bitar R, Leung G, Perng R, et al. MR pulse sequences: what every radiologist wants to know but is afraid to ask. Radiographics 2006;26:513 37. [3] Damadian R. Tumor detection by nuclear magnetic resonance. Science 1971;171:1151 3. [4] Elster AD. How much contrast is enough?. Dependence of enhancement on field strength and MR pulse sequence. Eur Radiol 1997;7(Suppl. 5): 276 80. [5] Gruber L, Loizides A, Luger AK, et al. Soft-tissue tumor contrast enhancement patterns: diagnostic value and comparison between ultrasound and MRI. AJR Am J Roentgenol 2017;208:393 401. [6] Tsalafoutas IA, Koukourakis GV. Patient dose considerations in computed tomography examinations. World J Radiol 2010;2:262 8. [7] Siriwanarangsun P, Statum S, Biswas R, Bae WC, Chung CB. Ultrashort time to echo magnetic resonance techniques for the musculoskeletal system. Quant Imaging Med Surg 2016;6:731 43. [8] FG S. MRI Safety; 2019. http://www.mrisafety.com/TMDL_list.php. [9] Bass E. Tendinopathy: why the difference between tendinitis and tendinosis matters. Int J Ther Massage Bodyw 2012;5:14 17. [10] Khan KM, Cook JL, Bonar F, Harcourt P, Astrom M. Histopathology of common tendinopathies. Update and implications for clinical management. Sports Med 1999;27:393 408. [11] Makino A, Pascual-Garrido C, Rolon A, Isola M, Muscolo DL. Mucoid degeneration of the anterior cruciate ligament: MRI, clinical, intraoperative, and histological findings. Knee Surg Sports Traumatol Arthrosc 2011;19:408 11. [12] Griffiths DL. Tenosynovitis and tendovaginitis. Br Med J 1952;1:645 7. [13] Delfaut EM, Demondion X, Bieganski A, Cotten H, Mestdagh H, Cotten A. The fibrocartilaginous sesamoid: a cause of size and signal variation in the normal distal posterior tibial tendon. Eur Radiol 2003;13:2642 9. [14] Erickson SJ, Cox IH, Hyde JS, Carrera GF, Strandt JA, Estkowski LD. Effect of tendon orientation on MR imaging signal intensity: a manifestation of the “magic angle” phenomenon. Radiology 1991;181:389 92.

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[15] Rosenberg ZS, Beltran J, Bencardino JT. From the RSNA refresher courses. Radiological Society of North America. MR imaging of the ankle and foot. Radiographics 2000; 20 Spec No:S153 179. [16] Fredericson M, Bergman AG, Hoffman KL, Dillingham MS. Tibial stress reaction in runners. Correlation of clinical symptoms and scintigraphy with a new magnetic resonance imaging grading system. Am J Sports Med 1995;23:472 81. [17] Bredella MA, Tirman PF, Peterfy CG, et al. Accuracy of T2-weighted fast spin-echo MR imaging with fat saturation in detecting cartilage defects in the knee: comparison with arthroscopy in 130 patients. AJR Am J Roentgenol 1999;172:1073 80. [18] Paunipagar BK, Rasalkar D. Imaging of articular cartilage. Indian J Radiol Imaging 2014;24:237 48. [19] Grassi A, Quaglia A, Canata GL, Zaffagnini S. An update on the grading of muscle injuries: a narrative review from clinical to comprehensive systems. Joints 2016;4:39 46. [20] Othman MI, Chew KM, Peh WC. Variants and pitfalls in MR imaging of foot and ankle injuries. SemMusculoskelet Radiol 2014;18:54 62. [21] Park HJ, Lee SY, Kim E, et al. Peroneal tendon pathology evaluation using the oblique sagittal plane in ankle MR imaging. Acta Radiol 2016;57:620 6. [22] Kellett JJ, Lovell GA, Eriksen DA, Sampson MJ. Diagnostic imaging of ankle syndesmosis injuries: a general review. J Med Imaging Radiat Oncol 2018;62:159 68. [23] Park HJ, Lee SY, Park NH, et al. Usefulness of the oblique coronal plane in ankle MRI of the calcaneofibular ligament. Clin Radiol 2015;70:416 23. [24] Weishaupt D, Treiber K, Kundert HP, et al. Morton neuroma: MR imaging in prone, supine, and upright weight-bearing body positions. Radiology 2003;226:849 56. [25] Hargreaves BA, Worters PW, Pauly KB, Pauly JM, Koch KM, Gold GE. Metal-induced artifacts in MRI. AJR Am J Roentgenol 2011;197:547 55. [26] Siriwanarangsun P, Bae WC, Statum S, Chung CB. Advanced MRI Techniques for the ankle. AJR Am J Roentgenol 2017;209:511 24. [27] Hopkinson WJ, Pierre St P, Ryan JB, Wheeler JH. Syndesmosis sprains of the ankle. Foot Ankle 1990;10:325 30. [28] Norkus SA, Floyd RT. The anatomy and mechanisms of syndesmotic ankle sprains. J Athl Train 2001;36:68 73. [29] Zanon G, Div G, Marullo M. Osteochondritis dissecans of the talus. Joints 2014;2:115 23. [30] Hildebrand KA, Frank CB. Scar formation and ligament healing. Can J Surg 1998;41:425 9. [31] Patten A, Pun WK. Spontaneous rupture of the tibialis anterior tendon: a case report and literature review. Foot Ankle Int 2000;21:697 700. [32] Schweitzer ME, Karasick D. MR imaging of disorders of the posterior tibialis tendon. AJR Am J Roentgenol 2000;175:627 35. [33] Hamilton WG. Stenosing tenosynovitis of the flexor hallucis longus tendon and posterior impingement upon the os trigonum in ballet dancers. Foot Ankle 1982;3:74 80. [34] Pedowitz D, Kirwan G. Achilles tendon ruptures. Curr Rev Musculoskelet Med 2013;6:285 93. [35] Thomas JL, Christensen JC, Kravitz SR, et al. The diagnosis and treatment of heel pain: a clinical practice guideline-revision 2010. J Foot Ankle Surg 2010;49:S1 19. [36] Theodorou DJ, Theodorou SJ, Farooki S, Kakitsubata Y, Resnick D. Disorders of the plantar aponeurosis: a spectrum of MR imaging findings. AJR Am J Roentgenol 2001;176:97 104. [37] Finkenstaedt T, Siriwanarangsun P, Statum S, et al. The calcaneal crescent in patients with and without plantar fasciitis: an ankle MRI study. AJR Am J Roentgenol 2018;211:1075 82. [38] Starr AM, Wessely MA, Albastaki U, Pierre-Jerome C, Kettner NW. Bone marrow edema: pathophysiology, differential diagnosis, and imaging. Acta Radiol 2008;49:771 86. [39] Crema MD, Roemer FW, Marra MD, et al. Contrast-enhanced MRI of subchondral cysts in patients with or at risk for knee osteoarthritis: the MOST study. Eur J Radiol 2010;75:e92 6. [40] Shakked RJ, Walters EE, O’Malley MJ. Tarsal navicular stress fractures. Curr Rev Musculoskelet Med 2017;10:122 30. [41] Zanetti M, Ledermann T, Zollinger H, Hodler J. Efficacy of MR imaging in patients suspected of having Morton’s neuroma. AJR Am J Roentgenol 1997;168:529 32. [42] Birt MC, Anderson DW, Bruce Toby E, Wang J. Osteomyelitis: recent advances in pathophysiology and therapeutic strategies. J Orthop 2017;14:45 52. [43] Allen LR, Flemming D, Sanders TG. Turf toe: ligamentous injury of the first metatarsophalangeal joint. Mil Med 2004; 169:xix xxiv. [44] Mugler 3rd JP. Optimized three-dimensional fast-spin-echo MRI. J Magn Reson Imaging 2014;39:745 67.

Chapter 20

Biomechanical Assessment of Soft Tissues in the Foot and Ankle Using Ultrasound Roozbeh Naemi1, David Allan1, Sara Behforootan2, Panagiotis Chatzistergos1 and Nachiappan Chockalingam1 1

Centre for Biomechanics and Rehabilitation Technologies, Staffordshire University, Stoke-on-Trent, United Kingdom, 2Imperial College London,

Department of Surgery and Cancer, London, United Kingdom

Abstract There has been an increasing use of ultrasound as an imaging modality in the assessment of foot and ankle function with implications in a range of diagnostic applications. The importance of the mechanical characteristics of the soft tissue within the foot and ankle and their relationship to physical and physiological functions during weightbearing activities have been previously established. Given this significance, this chapter has a focus on the application of ultrasound as an imaging modality for biomechanical assessment of the soft tissues of the foot and ankle. This chapter provides an overview of ultrasound-based assessment to examine the mechanical behavior of the soft tissues, expanding on methods like ultrasound assessment when the tissue is under realistic loading conditions and ultrasound elastography (sonoelastography). In addition, the role of ultrasound in assessment of the mechanical properties of plantar fat pad, plantar fascia, and Achilles tendon in healthy and pathological conditions are discussed.

20.1

Introduction

Ultrasound imaging is routinely used within medicine to aid in the diagnosis and treatment of numerous different conditions. As an imaging modality, ultrasound currently accounts for approximately 20% 25% of all medical scans performed, and this is increasing at a rate of approximately 5% per year [1]. This increase in the use of ultrasound, as a medical imaging tool, is due to its accessibility and its nonionizing characteristics. In some cases, it has replaced magnetic resonance imaging (MRI), computed tomography (CT), and radiography, and in many specific clinical settings it serves as an important additional tool in diagnosing clinical conditions. Ultrasound imaging has been utilized in a range of lower extremity diagnostic and interventional applications. There has been an increasing use of ultrasound as an imaging modality in foot and ankle assessments, as most anatomical structures are superficial, discrete, and have clear landmarks [2,3]. The applications include ultrasound examination ranging from the diagnosis of mass lesions to identification of nerve entrapment syndromes. It is also common for ultrasound to be used during interventions needing accurate needle placement [2]. Given the importance of the mechanical characteristics of the soft tissue within the foot and ankle and their relationship to physical and physiological function during weightbearing activities, this chapter focuses on the application of ultrasound imaging for biomechanical assessment of the soft tissues around the foot and ankle.

20.1.1 Ultrasound imaging, how it works Ultrasound imaging works by using a transducer to produce an ultrasound wave with a frequency usually ranging from 2 18 MHz [4]. The ultrasound waves produced by the transducer are transmitted into the tissues and are reflected back Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00022-6 © 2023 Elsevier Inc. All rights reserved.

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by the different tissues to varying degrees. Using the speed of sound in tissues and the time of each waves return, the ultrasound scanning unit calculates the distance from the transducer to the boundaries between tissues, as well as to small irregularities within tissues. These distances are then used to generate two-dimensional images of tissues and organs (B-mode). The quality of the ultrasound image and the penetration depth of ultrasound waves depend on the transducer frequency being used. The lower the frequency of the ultrasound wave, the greater the penetration into the soft tissues, allowing the deeper tissues and organs of the body to be imaged. The higher the transducer frequency, the lower the penetration depth, but the greater the resolution of the image, resulting in a clearer picture. The amount of echo returned after hitting a tissue interface is determined by the difference between the two tissues’ acoustic impedance. Acoustic impedance is proportionate to the density of the tissue and the US wave propagation velocity in the tissue. Hence no echo is generated at the interface between two tissues with identical acoustic impedance. While the interfaces between soft tissues of similar acoustic impedances usually generate lowintensity echoes, strong echoes are generated at the interfaces between soft tissue and bone [5]. The amount of echo detected by the transducer also depends on the relative angle between the tissues’ interfaces and the ultrasound beam (direction of traveling waves), with the maximum when the interface is perpendicular to the US beam. Depending on the amount of echoes detected the hyperechoic areas usually appear as bright while the hypoechoic areas appear as dark in the B mode ultrasound. Also, in extreme cases anechoic refers to areas with minimal echo detected and appear in black, and with contrast echogenic refers to the areas with maximum echo detected and appear as bright white. As the speeds of sound transmission is different in two different tissues, a change in the direction of waves known as refraction occurs when the sound hits the interface between two tissues. As the sound frequency is constant, to accommodate the different speed of sound transmission, the wavelength has to change in the two tissues. This results in a redirection of the sound pulse as it passes through the interface. The speckle signal that provides the visible texture in tissues like muscle is a result of interface between multiple scattered echoes produced by the scattering structures within the imaged plane [6]. Sonoelastography (sometimes referred to as ultrasound elastography) is an ultrasound imaging technique in which the stiffness of soft tissues can be quantified and visualized. The term elastography was first described by Ophir et al. as a method of measuring and visualizing the strain properties of biological tissues [7]. In this method, the tissue is deformed and measured either to calculate and display strain (strain elastography) or to calculate and display shear wave speed (shear wave elastography). The tissue displacement can be in the form of either normal deformation that is usually produced by using the ultrasound transducer to deform the tissue surface or in the form of shear deformation that is usually produced by acoustic radiation force [8]. In strain elastography static, quasistatic, or dynamic forces can be used, while the shear wave elastography requires the creation of a shear wave, which in turn requires the use of a dynamic force, albeit with a much smaller magnitude [8]. In strain elastography the transducer moves axially upwards and downwards which causes the tissue to be pressed slightly at a low pace (i.e., 1 2 Hz). The strain caused in the tissue is defined as a function of correlation (similarity) between subsequent ultrasound images of the deformed tissue. The extent of similarity between pre and postcompression ultrasound echoes are usually quantified based on the cross-correlation method. The relative average values of strain are inversely correlated with tissue stiffness [9]. As an indicator, the software used within strain elastography machines (Fig. 20.1) provides real-time feedback to the operator to maintain the quality of compression, which may require a low compression rate (i.e., 1 2 Hz) and low deformation (i.e., 1 2 mm). However, due to the hyperelastic characteristics of the soft tissue, the force-deformation behavior is highly nonlinear, and the strainability measured is highly dependent on the amount of force applied. To decrease the effect of variations in loading on evaluations of stiffness, the force applied can be regulated. This can be achieved by using an interface material (such as a standoffs with known constant stiffness) during scans that are made from soft deformable materials that allow the passage of ultrasonic waves without producing any echoes from reflections. The standoff can be used between the probe and the skin during strain elastography, where the ratio of the stiffness of the tissue of interest is calculated based on the deformation of tissue relative to the amount of deformation of the standoff material, former representing an estimation of applied force (Fig. 20.1) [10]. Shear wave elastography is another noninvasive, ultrasound-based method for the assessment of the soft tissues’ stiffness. It involves the generation of shear waves inside the tissue and the measurement of their propagation speed as they expand laterally in the field of view. In shear wave elastography, an acoustic radiation force impulse (a focused ultrasound pulse produced by the ultrasound transducer) deforms the tissue perpendicular to the transducer, and a shear wave is produced traveling in parallel to the transducer.

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FIGURE 20.1 The B-mode ultrasound image of heel pad and stand-off under maximum compression (left). The elastography image superimposed over the same B-mode image (right). The ellipses used for measuring relative stiffness and thickness are also shown. In all cases the major axis of ellipse is equal to the width of the image (22.7 mm), with its minor axis adjusted to the thickness of the soft tissue and standoff respectively for zone 1 and 2. The area of each zone as the area of ellipses for zone 1 (A Z1) and zone 2 (A Z2), together with the relative strainability of zone 2 (standoff) to zone 1 (tissue) appears at the left end of the screen (ELX2/1). From Naemi R, Chatzistergos P, Sundar L, Chockalingam N, Ramachandran A. Differences in the mechanical characteristics of plantar soft tissue between ulcerated and non-ulcerated foot. J Diabetes Complicat 2016;30 (7):1293 9. ,http://www.sciencedirect.com/science/article/pii/S1056872716301830. [accessed 29.09.16].

As the shear waves expand perpendicular to the ultrasound pulse (i.e., parallel to the transducer), the speed of shear wave propagation is measured. The speed of the shear wave is used to estimate the tissues shear modulus (G) and Young’s modulus (E) using the following equation: E 5 3G  3ρC2

(20.1)

Where C is the recorded shear wave speed for the area of soft tissue and ρ is the density of the selected tissue (ρ  1000kg=m3 for soft biological tissues). The stiffness of different soft tissues can then be estimated based on the measured shear wave speed, assuming an isotropic solid material (Fig. 20.2).

20.2 Ultrasound assessment of structural changes: the effect of weightbearing activities Ultrasound is a reliable method to quantify the structural integrity of the foot muscles and plantar fascia in a nonweightbearing condition [11]. While changes in the structure, that is, in the foot muscle cross-sectional area, in weightbearing conditions have been reported compared to nonweightbearing, measurements in weightbearing conditions have also been reported to be reliable [12].

20.2.1 Assessment of plantar soft tissue thickness in relation to weightbearing activities Several studies designed custom apparatuses to allow weightbearing ultrasound imaging of the plantar soft tissue [13 15]. These devices generally consisted of a frame that supports the foot during standing/walking, which allows the probe to be positioned so that the plantar soft tissues, that is the heel pad and submetatarsal fat pad, can be pushed against the transducer and the loaded tissue thickness can be measured.

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FIGURE 20.2 A typical SW elastography image of the heel pad under minimum compression. The interface between the more superficial, stiffer layer 1 and the deeper, softer layer 2 is highlighted with a horizontal dotted line. This interface was used to define the cyclic area where SW speed was assessed and to measure the deformation of layer 1. From Chatzistergos PE, Behforootan S, Allan D, Naemi R, Chockalingam N. Shear wave elastography can assess the in-vivo nonlinear mechanical behavior of heel-pad. J Biomech 2018. ,https://www.sciencedirect.com/science/article/pii/ S0021929018307243?dgcid 5 rss_sd_all. [accessed 7.10.18].

The scanning technique proposed by Bygrave and Betts [15] was utilized in a study by Young and coworkers where plantar pressure was found to be negatively correlated to plantar soft tissue thickness at metatarsal heads [16]. The weightbearing ultrasound measurement technique was deemed to be useful in the screening of diabetic and rheumatoid patients for high metatarsal pressures where plantar pressure measuring systems are not available [16]. The association between peak plantar pressure during walking and plantar soft tissue thickness measured in standing was later confirmed in a study of diabetic neuropathic patients [17]. Another study reported a significant difference in the heel pad thickness between weightbearing and nonweight bearing conditions and postulated that a weightbearing scan may allow more objective assessment of plantar soft tissue pathologies [13]. Using a similar device in a later study, this group demonstrated a significant difference in heel pad thickness between runners with heel pain compared to runners without heel pain [18]. Ultrasound has also been used in the assessment of plantar soft tissue thickness during the stance phase of walking. In a study by Cavanagh on a group of healthy participants, an average strain of 45.7% for the plantar soft tissue underneath the second metatarsal head was reported during the late stages of ground contact during barefoot walking [14]. To assess the strain of the plantar soft tissue in a shod condition, ultrasound was used within a shoe to assess the strain of the plantar heel pad during walking [19] (Fig. 20.3). The results when used in combination with motion capture indicated to have implications in providing more accurate measurements of the intrinsic foot kinematics by reducing soft tissue artifacts [20]. Another study using a similar methodology, showed that a custom molded silicone insole was able to reduce soft tissue strain more effectively when compared to a prefabricated insole [21]. This study also suggested that quantifying the reduction of soft tissue strain, in addition to reduction in the plantar pressure, could be considered as an essential design requirement for orthotic insoles. The authors claimed that this would allow an effective design of an insole to prevent ulcer initiation in people with diabetes [21].

20.2.2 Assessment of plantar fascia thickness in relation to weightbearing activities The reliability of ultrasound in measuring plantar fascia thickness in nonweightbearing conditions has been established [22], and the relationship to foot arch shape and plantar pressure during gait in individuals with unilateral heel pain was investigated [23]. This study reported a thicker plantar fascia in symptomatic limbs compared to asymptomatic, while fascial thickness, in turn, was positively correlated with arch angle in both symptomatic and asymptomatic feet, and with pain and peak regional loading of the midfoot in the symptomatic limb [23]. The measurement of plantar fascia thickness during weightbearing conditions was reported to be higher in both the symptomatic and asymptomatic feet of patients with plantar heel pain compared to the control healthy group [24]. It was concluded that ultrasonic scanning, when incorporated in a platform as proposed earlier [15], can improve the

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(A) (B)

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FIGURE 20.3 In-shoe embedded US setup. Left top: CAD design for inserts, (A) heel cup insert; (B) flat insert. Left bottom: US transducer. Right: modified training shoe fitted with embedded transducer. From Telfer S, Woodburn J, Turner DE. Measurement of functional heel pad behavior inshoe during gait using orthotic embedded ultrasonography. Gait Posture 2014;39(1):328 32. ,http://www.ncbi.nlm.nih.gov/pubmed/23962596. [accessed 12.10.18].

standardization and reproducibility of measurement [24]. For further details on the application of ultrasound assessment of plantar fascia in nonweightbearing conditions for diagnosing plantar fasciitis, the readers are encouraged to consult the two systematic reviews of the literature in this area [25,26].

20.2.3 Assessment of Achilles tendon thickness in relation to weightbearing activities Ultrasound imaging has been used in numerous studies to measure the changes in the thickness of Achilles tendon during weightbearing activities. In a study assessing the effect of walking on Achilles tendon strain, changes in the diameter of the tendon were measured at the start and end of the day, and a positive significant relationship was reported between the amount of walking activity and reduction in tendon diameter [27]. Another study reported that in contrast to preexercise measures, the Achilles tendon thickness measured immediately following conditioning exercise, against an effective resistance between 100% and 150% body weight, was significantly correlated with body weight and body mass index (BMI) [28]. To further assess the effect of BMI on Achilles tendon mechanical characteristics, another study explored the use of longitudinal sonograms of the Achilles tendon prior to and following weight-bearing ankle exercises to measure the reduction in transverse tendon thickness for normal and overweight participants [29]. While the Achilles tendon was reported to be thicker in the people who are overweight when compared to normal group both prior to and immediately after exercise, the immediate changes in the thickness of Achilles tendon as a result of exercise in the overweight group (-10.7 6 2.5%) was almost half compared to the corresponding values in the normal-weight group (-19.5 6 7.4%) [29]. The authors linked these findings to the effect of obesity on the structural changes in tendon that impairs intratendinous fluid movement in response to load [29]. The changes in the sonographic characteristics and anteroposterior thickness of Achilles tendon was also investigated in response to eccentric exercise in a group of participants with unilateral Achilles tendinopathy and in healthy controls [30]. The results showed that compared to the tendon in the control group, both the asymptomatic and symptomatic tendons were significantly thicker and hypoechoic before exercise, while in all participants the tendon’s thickness significantly decreased immediately after eccentric exercise [30]. The symptomatic tendon was characterized by a significantly lower reduction in anterior/posterior thickness immediately following eccentric exercise compared to both the asymptomatic tendon and the tendons in the control group, where there were no changes in the anterior/posterior thickness as an immediate effect of exercise [30].

20.2.4 Summary and limitations of weightbearing ultrasound Overall, B-mode ultrasound seems to be a useful tool to reliably assess the immediate changes in the thickness of the soft tissues of the foot, namely the plantar fat pad, plantar fascia, and Achilles tendon, in response to or as a result of weightbearing activities. While this can provide important information on the mechanical behavior, that is, the changes in dimension of tissue when under load or as an immediate result of loading, to understand the mechanical properties

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(i.e., stiffness or modulus) of the soft tissue in relevant conditions, the deformation under known loads needs to be measured.

20.3

Ultrasound assessment combined with measurement of load

Ultrasound imaging has been used in conjunction with force/torque measurements to assess the mechanical properties of soft tissues in the foot and ankle. As soft tissues exhibit visco-hyperelastic characteristics in which the forcedeformation relationship is highly nonlinear and affected by deformation rate, specific considerations need to be given to ensure a reliable measurement of the soft tissue characteristics using these methods.

20.3.1 Ultrasound combined with load cells to assess the mechanical properties of the plantar soft tissue The mechanical properties of the plantar fat pad is an important area of study due to its role in foot-ground contact during weightbearing activities of daily living [31]. The heel pad during normal gait provides shock absorption and distributes compressive loading. There is an increasing body of evidence indicating that pathological conditions such as plantar heel pain or diabetic foot complications are linked with changes in the force-deformation relationship of plantar soft tissue [32 37]. Hence ultrasound imaging has been used to assess the in vivo mechanical behavior of the heel pad during loading, in a method commonly known as ultrasound indentation. In this type of test, using a load cell and an ultrasound probe in series, a simultaneous measurement of applied force and the deformations of soft tissue is conducted and used to assess the stiffness of the soft tissue [10,33,37 44]. The ultrasound probe indents the foot while the load cell measures the load instantaneously and the deformation of the soft tissue is measured generally using the Bmode ultrasound images (Fig. 20.4) [41]. In a study using an ultrasound indentation device it was reported that triglycerides and blood sugar levels are significantly correlated with the stiffness of the heel pad in diabetic patients [41]. While this study found associations between the mechanical properties of the heel pad and the clinical parameters that could be linked to foot ulceration risk, it concluded that changes in heel pad stiffness could limit the tissues’ ability to evenly distribute loads making them more vulnerable to trauma and ulceration [41]. Parametric models that represent the force-deformation behavior of the heel pad have been used to quantify the nonlinear viscoelastic characteristics of the plantar soft tissue; for a summary, see elsewhere [31]. More recently, Naemi and colleagues proposed a method that can differentiate between the elastic and viscous components of the reaction force that contribute to the total stiffness of the heel pad [45]. Using data from ultrasound indentation these authors used the difference between the loading and unloading curves of the force-deformation data to assess the viscous and elastic components and to quantify how these two components change at different deformations and deformation rates using reaction force model parameters [45].

FIGURE 20.4 A custom made device used for in vivo loading tests: (a) ultrasound probe, (b) dynamometer, (c) probe holder, (d) ball-screw actuator, (e) hand wheel and (f) foot support. From Chatzistergos PE, Naemi R, Sundar L, Ramachandran A, Chockalingam N. The relationship between the mechanical properties of heel-pad and common clinical measures associated with foot ulcers in patients with diabetes. J Diabetes Complicat 2014;28 (4):488 93.

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The results from this study also indicate that the parameters that show viscosity were significantly correlated with maximum strain, indicating the need to perform indentation tests at strain rates relevant to walking to assess the mechanical behavior of plantar soft tissue in realistic conditions [45]. This concept was further developed in another study where individual-specific mathematical models extracted from ultrasound indentation of heel pad at realistic deformation rates, combined with the plantar pressure measurements during gait, were used to calculate the heel pad strain during walking [40]. In this study it was found that the calculated heel pad strain during walking can be predicted to within 14% of the in vivo values that were measured using motion capture [40]. While no statistically significant correlation was observed between maximum strain and peak plantar pressure during walking; the study indicated that the measurement of strain along with plantar pressure can clarify the risk of mechanical trauma to the plantar soft tissue [40]. Ultrasound indentation tests have also been used in conjunction with subject-specific geometry gathered from ultrasound to develop finite element models and to reverse engineer the stress-strain graph of the tissue, to find geometryindependent mechanical properties of soft tissue [38,39,42,46]. For further information the readers are encouraged to read the finite element chapter in this book. In all of these studies, the indentation of the heel pad was conducted with the participant positioned in a supine position on an examination table. However, few studies have measured and quantified the mechanical properties of the foot when a participant is in an upright position [44,47,48]. Indentation in an upright weightbearing condition allows the effect of preloading (i.e., during standing and walking) on the mechanical properties of the plantar soft tissue to be investigated. For example, testing healthy participants with loading ranging from 0% to 80% of bodyweight, the heel pad thickness was found to decrease by 12.0% while the stiffness increased by 83.4% [48]. The device was also reported to be able to implement complex loading patterns similar to gait [44] and was used to calculate the material properties of the heel pad sublayers using an inverse finite element approach [47]. Although ultrasound indentation devices that are utilized for subjects in standing or supine positions can produce compression rates to replicate the deformation of the heel pad during walking, the exact replication of this behavior would be challenging and require a force-control mechanism. Another study embedded an ultrasound probe into an orthotic device placed within a standard sneaker to investigate the force-deformation behavior of the heel pad during walking [19]. This method was shown to be reliable to measure person-specific stiffness and energy dissipation ratio of the heel pad during gait and showed the efficacy of a contoured heel cup in reducing the compression of the mid portion of the heel pad during walking [19].

20.3.2 Ultrasound combined with dynamometry to assess the mechanical properties of Achilles tendon Ultrasound has been combined with a dynamometer to calculate the mechanical properties of the Achilles tendon. This approach generally combines the use of B-mode ultrasound imaging and an isokinetic dynamometer to measure the torque being produced around the ankle [49 55]. Within these studies the Achilles tendon is imaged at rest and during maximal or sub maximal plantar/dorsiflexion to calculate the elongation within the tendon (Fig. 20.5) [54]. The simultaneous measurement of torque is then divided by the length of the tendon moment arm, to give a value for the force by which the tendon is stretched. Combining these measurements allows the extraction of the force deformation behavior of the tendon and the calculation of its stiffness. The study of Achilles tendon stiffness at maximum isometric voluntary contraction was reported to be in agreement with the preexisting data based on in vitro testing of isolated tendons. It was suggested that under physiological loading the Achilles tendon operates within the elastic “toe” region [53]. The method described in this study series [53,56,57] has since been used in several subsequent studies including the investigation of: tendon-aponeurosis strains [49]; the differences in tendon stiffness in the dominant versus nondominant leg [50]; the effect of loading history on tendon stiffness [55], and the relationship between tendon stiffness and race performance [52]. In addition using a similar methodology the mechanical properties of the Achilles tendon has been investigated in tendinopathy [51]. In a study using ultrasound combined with dynamometry a significantly higher Young’s modulus and length of Achilles tendon in the dominant compared to the nondominant leg was reported. This was attributed to the differences in the loading profiles between legs during daily activities in the healthy population [50]. Using ultrasound dynamometry, another study reported the Achilles tendon stiffness in highly trained male long distance runners to be negatively correlated to race performance [52]. Achilles tendon stiffness was also reported to be unchanged between the pre and post marathon [55]. It was concluded that the Achilles tendon of the physically active individuals seems to be able to resist mechanical changes under physiological stress suggesting that activities such as running may not predispose the tendon to injury [55].

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FIGURE 20.5 Example of sequential US images that show the myotendinous junction displacement (arrow) at plantarflexion, neutral and dorsiflexion. From de Oliveira LF, Peixinho CC, Silva GA, Menegaldo LL. In vivo passive mechanical properties estimation of Achilles tendon using ultrasound. J. Biomech 2016;49(4):507 13.

Ultrasound dynamometry has also been used to investigate the mechanical properties of Achilles tendon in tendinopathy. In a study on a pathological population, the tendon characteristics (shape, composition) and mechanical properties (strain, stiffness) of the involved side of participants with insertional Achilles tendinopathy (IAT) were compared to the uninvolved side and to controls [51]. This study reported a significantly larger tendon diameter, lower echogenicity, higher strain, and lower stiffness in the side with IAT when compared to the side without and to healthy controls [51]. In most studies using ultrasound dynamometry, the strain of the Achilles tendon was measured in the longitudinal direction. However, several studies also reported strain in the transverse direction [54,58,59]. The Poisson’s ratio of Achilles tendon, indicating the degree at which the tendon shrinks in width when stretched, was measured using ultrasound and was reported to be highly individualized [54]. 3D free hand ultrasound was also used to study the deformation of the Achilles tendon, with significantly greater longitudinal strain (5.2% 6 1.7%) reported for the whole tendon compared to the proximal Achilles tendon (2.6% 6 2.0%) during submaximal fixed-end contractions of the triceps surae [58]. In this study a considerable transverse (orthogonal to the longitudinal direction) strain (5.0% 6 4%) was also reported for the proximal Achilles tendon during contraction [58]. In another study using 3D free hand ultrasound of the Achilles tendon at 70% maximum voluntary contraction of the plantarflexors, a significant increase in anteroposterior diameter (8.7%); and significant reductions in cross-sectional area (-5.5%) and mediolateral diameter (-8.7%) (all average values over length of the tendon) were reported relative to the corresponding values in resting condition [59]. An increased transverse rotation of the tendon was also reported during contraction when compared to rest, resulting in the tendon becoming externally rotated relative to the calcaneal

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insertion, with the peak strain happening at the midportion of the Achilles tendon. The method was deemed to have important implications in estimating the tendon stress in vivo [59]. To investigate the mechanical behavior of the Achilles tendon during weightbearing activities, ultrasound measurements of Achilles tendon were conducted during a hopping exercise where motion capture and force plates were utilized to approximate the Achilles tendon load using an inverse dynamics approach [60]. This study reported an average energy return efficiency (the percentage of strain energy that is return during unloading) of 16%, where the average peak strain was reported to be 8.3% during the hop, and concluded that prolonged hopping may cause tendon damage due to the observed excess strain [60].

20.3.3 Summary and limitations of ultrasound assessment combined with measurement of load Indentation of heel pad using ultrasound probe can quantify the bulk stiffness of heel pad under controlled conditions. However, to be able to assess realistic person-specific force-deformation behavior the tests need to replicate the loading rate and magnitude at which the heel pad deforms during weightbearing activities. As also mentioned earlier, the stiffness measured during dynamic loading includes both a viscous and an elastic component, out of which the viscous component is dependent on the deformation rate at which the plantar soft tissue is compressed, while the elastic component is only dependent on the amount of deformation [45]. This indicates that the higher the deformation rate of the tissue, the higher the stiffness. Hence to assess the mechanical properties of soft tissue under realistic conditions, the deformation rate and magnitude during indentation tests should be similar to those experienced during walking. Often times, for both technical and safety reasons slower rates and lesser magnitudes are employed. Ultrasound combined with dynamometry has been well established and widely used in biomechanical studies, but it is understood that the measures of Achilles tendon stiffness can be influenced by the coactivity of the tibialis anterior during isometric contraction. This can have an influence on the calculation of the force applied to the Achilles tendon and hence on the measurement of stiffness [49]. For further information the readers are encouraged to consult the systematic review and the critical overview of the literature surrounding the ultrasound-based testing of tendon mechanical properties [61,62]. Overall, the method proposed in this section, where ultrasound imaging is combined with direct measurement of force on the tissue, can give a bulk assessment of the mechanical properties of the soft tissue including strain or stiffness. However, the other important aspect that is the regional mechanical properties cannot be assessed with these methods. In the following section, ultrasound elastography techniques that can provide a detailed map of the soft tissues’ mechanical properties will be discussed. Another limitation of the current method of ultrasound indentation is that it only focuses on axial stiffness of the soft tissue by simultaneously measuring axial deformation and force along the same direction. However, in addition to the compressive properties of the heel pad, the shear properties are also important in assessing the vulnerability of tissue to mechanical trauma and hence there is a need for ultrasound dynamometry techniques that can assess the shear properties of the heel pad as well [31].

20.4

Ultrasound elastography (sonoelastography)

In the following subsection, the applications of elastography in the assessment of mechanical properties of the soft tissues in foot and ankle, including the heel pad, plantar fascia, and Achilles tendon, are discussed.

20.4.1 Ultrasound elastography to assess the mechanical properties of plantar soft tissue Shear wave elastography has been used to study the mechanical properties of the different soft tissues of the heel pad in several studies [63 65]. It was used in a study of healthy participants, where the mechanical properties of the heel pad were found to be highly heterogeneous, with the stiffness of the heel found to be greatest beneath the plantar skin and continuously decreasing through the micro- and then macro-chambers [63] (Fig. 20.6). A similar methodology was also used to study the heel pad stiffness of participants with unilateral plantar heel pain, where the stiffnesses of micro- and macro-chambers were found to be significantly higher in the involved foot compared to uninvolved foot [64]. Shear wave ultrasound has also been used in another study where the stiffness of the lateral soft tissue of the heel was also explored. It was found that prolonged loading (over 90 and 150 min) lead to softening of deep subcutaneous structures and stiffening of the skin [65] and this was considered to play a key role in pressure ulcer development. Shear wave elastography was postulated to serve as a tool for quantifying the risk of developing pressure ulcer damage [65]. Ultrasound strain elastography has also been used in studying the mechanical properties of the heel pad and first submetatarsal fat pad in patients with diabetic foot ulceration [10]. A strain elastography technique together with the

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FIGURE 20.6 SWE measurement. The measurement was conducted using three 10 mm long measurement lines along the depth direction of the heel pad. The maximum measurement depth was set as 10 mm because SWE signals may not always be able to penetrate over 10 mm, probably owing to ultrasound attenuation, although in some cases signals could penetrate the whole heel pad thickness. The calcaneal tubercle is marked with an arrow. From Lin C-Y, Chen P-Y, Shau Y-W, Tai H-C, Wang C-L. Spatial-dependent mechanical properties of the heel pad by shear wave elastography. J Biomech 2017;53:191 195.

use of a stand-off material (Fig. 20.1) as reference enabled a quantitative assessment of stiffness, and a lower heel pad stiffness was observed in patients with active foot ulcers [10]. Although the results reported by Naemi and coworkers could indicate a possible link between tissue mechanics and ulceration [10], this did not identify whether the observed differences were due to the physiological changes that contribute to ulceration or whether those reflected the pathophysiological changes that happen after ulceration. To address this issue, in a prospective study using the same measurement methodology, a higher plantar soft tissue thickness and lower ratio of stiffness to thickness at the first submetatarsal fat pad in patients with subsequent ulceration incident (during a year followup) were reported [66]. It was also reported that the inclusion of the mechanical properties of plantar soft tissue in predicting diabetic foot ulceration improved the sensitivity and prognosis accuracy for ulceration by 14% and 5% respectively [66].

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20.4.2 Ultrasound elastography to assess the mechanical properties of plantar fascia Sonoelastography has been utilized in recent years to investigate the effect of plantar fasciitis on the stiffness of the plantar fascia [67 69], to assess the diagnostic accuracy of plantar fasciitis [70,71], and to investigate the treatment efficiency of plantar fasciitis [72]. Using sonoelastography, Wu and coworkers revealed that the plantar fascia stiffness correlated negatively with age and was reported to be lower in patients with plantar fasciitis [67]. This was supported by results from a later study using sonoelastography, where the plantar fascia stiffness of patients who were at the early diagnosis stage of plantar fasciitis was found to be significantly lower compared to the control group [69]. The plantar fascia in patients with typical clinical manifestations of plantar fasciitis was reported to be less stiff in either asymptomatic or symptomatic sides of unilateral plantar fasciitis patients, and both sides of bilateral plantar fasciitis patients when compared to healthy controls [68]. Furthermore, the pain score in patients with plantar fasciitis was significantly correlated with lower stiffness of the plantar fascia [71]. The diagnostic accuracy of real-time sonoelastography in plantar fascia was compared to that of a standard ultrasound when MRI was used as a gold standard for diagnosing plantar fasciitis [70]. In this study, a diagnostic accuracy of 96% was reported for real-time sonoelastography. This was much higher than standard B-mode ultrasound that showed a diagnostic accuracy of 68% [70]. Similarly, in another study, a high interobserver reproducibility was reported using strain sonoelastography. It was found that the overall accuracy increased significantly from 90.0% with B-mode ultrasound to 95.4% with sonoelastography [71]. While these results are comparable to that of MRI, ultrasound elastography has the advantage compared to MRI of being a quicker and less costly option. Sonoelastography was also used to evaluate the effectiveness of collagen injections in the treatment of plantar fasciitis [72]. The results indicated that after treatment, the plantar fasciitis was significantly stiffer, highlighting a potential clinical application of real-time sonoelastography which could aid in monitoring the treatment of plantar fasciitis [72].

20.4.3 Ultrasound elastography to assess the mechanical properties of Achilles tendon Sonoelastography has also been used to study the mechanical properties of Achilles tendon in healthy participants [73,74]. Using shear wave elastography, the effect of different levels of dorsiflexion was observed on the axial and sagittal stiffness of the Achilles tendon [73]. While a significant increase in tendon stiffness and anisotropy as a result of dorsiflexion were observed, interobserver correlation coefficients of 0.43 and 0.46 for the evaluation of Young’s modulus and shear wave speed respectively, were reported [73]. Ultrasound elastography has a moderate [74] to high [75] inter-operator measurement reliability for the assessments of Achilles tendon stiffness. The intra-operator reliability was also reported to be moderate [75] to high [74,76]. Sonoelastography was also used to assess the differences between Achilles tendon stiffness in people who exercise frequently and those who do not [74]. This study reported a significantly stiffer Achilles tendon in the nondominant foot in people who exercise frequently compared to people who did not, whereas no significant difference in Achilles tendon stiffness for the dominant foot between the two groups was reported [74]. Shear wave elastography has also been used to assess the stiffness of Achilles tendon in pathological conditions such as chronic overuse-associated pain [77] and in tendinopathy [78,79]. The Achilles tendon of patients with chronic overuse-associated pain was found to be stiffer and thicker compared to the asymptomatic volunteers [77]. However no such difference between symptomatic and control tendons at the enthesis and myotendinous junction was reported [77]. Shear wave elastography has also been used for the assessment of Achilles tendon stiffness in relaxed position for individuals with Achilles tendinopathy, where a significantly lower stiffness in the axial (but not in the sagittal) direction for tendons with tendinopathy (compared to normal tendons) was reported [78]. However, in stretched positions significantly lower stiffnesses both in sagittal and axial directions were reported for tendons with tendinopathy compared to normal tendons [78]. It was also reported that the sensitivity of the method for diagnosing Achilles tendinopathy was relatively lower in relaxed versus stretch positions and higher in axial versus transverse directions, with the axial stiffness in stretched position showing highest sensitivity (66.7%) and the axial stiffness in relaxed position showing the highest specificity (91.5%) [78]. This could indirectly indicate that assessment of the tendon in loaded conditions reveals changes in the stiffness that may not be revealed when the tendon is scanned in a relaxed position, hence can indicate the importance of assessment in relevant loading conditions. The diagnostic accuracy of axial sonoelastography in confirming clinically diagnosed Achilles tendinopathy was compared against the B-mode ultrasound [79]. The diagnostic accuracy of sonoelastography (97.8%) was found to be higher than the B-mode ultrasound (diagnostic accuracy 94.7%) in confirming clinically symptomatic Achilles

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tendinopathy [79]. While the study found a moderate correlation between strain ratio and the clinical diagnosis scores for Achilles tendon, a significantly higher strain ratio was reported in the clinically diagnosed tendinopathy group compared to the healthy group [79]. Sonoelastography was also used to assess histologic degeneration in cadaveric Achilles tendons [80] and in diabetic patients [81]. When comparing sonoelastography with B-mode ultrasound to assess degeneration in cadaveric Achilles tendon, both methods verified all histologically normal tendons as normal indicating 100% specificity [80]. However, in identifying tendons with histologic degeneration, sonoelastography was found to have more sensitivity (100%) compared to B-mode ultrasound (86% sensitivity) [80]. In another study, sonoelastography revealed stiffer and thicker Achilles tendon in diabetic patients with ulcers when compared to those without foot ulcers [81]. In the group with no foot ulceration, Achilles tendon thickness was found to be positively correlated with neuropathy, retinopathy, nephropathy, peripheral arterial disease, and coronary arterial disease, while no significant correlation was detected in the group with ulceration [81]. Ultrasound elastography has also been used to assess the stiffness in healing Achilles tendons after surgical repair of a tendon rupture [82] or for comparison of stiffness of tendon in patients with surgically repaired complete ruptures [83]. Tendon function, using the American Orthopedic Foot and Ankle Society rating system of the repaired Achilles tendon, was evaluated at 12, 24, and 48 weeks postoperatively and found to be positively correlated with the elasticity of the repaired Achilles tendon [82]. This study proposed that shear wave elastography can provide biomechanical information for evaluating the mechanical properties of healing Achilles tendon and predict Achilles tendon function [82]. In another study the surgically repaired complete ruptures and their contralateral asymptomatic Achilles tendons were assessed with ultrasound and sonoelastography [83]. In this study, the ruptured tendons were found to have a less homogeneous stiffness compared to the healthy Achilles tendons [83]. For further information on the use of ultrasound elastography in assessment of Achilles tendon injuries the readers are encouraged to consult the referenced systematic review of literature in this area [84].

20.4.4 Summary and limitations of ultrasound elastography Ultrasound elastography is a versatile tool that allows relative tissue stiffness to be quantified without the need to measure load, which makes it feasible to use in a clinical setting. However, there are some limitations regarding the use of strain and shear wave elastography within clinical and research practice. As explained earlier, one of the main limitations of strain elastography is the variability in the pressure applied to the tissue adversely affecting the results. Another issue with the use of strain elastography is that the sonograms displayed in some machines in which the absolute stiffnesses are scaled to represent a relative deformability based on the range of deformability of tissue in the region of interest. This does not allow a direct absolute quantitative measure to be taken from these elastograms, as the stiffness of each individual structure within the tissue is dependent on the range of stiffness that can be detected in the region of interest. To account for these limitations of using ultrasound strain elastography, a stand-off material could be used between the ultrasound probe and the tissue. Stand-offs are made from soft deformable materials that allow the passage of ultrasonic waves without producing any echoes from reflections. Comparing the relative deformability of different tissues to that of the stand-off produces a quantitative assessment of absolute stiffness taking into account the differences in the applied load and inter-trial variations in the displayed stiffnesses [10]. Shear wave elastography in general can provide a measure of stiffness of various soft tissues, by measuring the propagation speed of shear waves through tissues. However, similar to strain elastography in some machines these shear wave speeds (and the calculated shear and Young’s moduli) are displayed as a color-coded image super imposed on the B-mode, with the colors adjusted to present a continuous spectrum ranging between the highest stiffness (i.e., in red) and the lowest stiffness (i.e., in blue) within the region of interest. This limitation does not allow a comparison of absolute soft tissues stiffnesses, and was highlighted by several studies [67 72,74,85]. To overcome this issue, Yamamoto and coworkers used an interface referred to as an “acoustic coupler” with a known Young’s modulus that allowed the absolute stiffness values of the tendon to be calculated as a proportion of the strain of the acoustic coupler [75]. Several studies also reported the shear wave speed [73] or used the shear wave speed to calculate an absolute stiffness value of the Achilles tendon [74,82]. While this allowed the absolute stiffness of the soft tissue to be calculated, care should be taken when comparing the results from one machine to another as the type of shear wave propagated in the tissue may be different in different machines, for example, conical versus conventional [8]. This is further elaborated in one study where a linear relationship was reported between the shear wave-based and finite element-based estimations of nonlinear mechanical behavior of the heel pad [86].

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A positive significant relationship between shear wave speed and the magnitudes of compression, where the study reported that shear wave elastography is capable of reliably assessing differences in stiffnesses, but the absolute values of stiffness should be used with caution [86].

20.5

Conclusion and future areas of research

This chapter provides a review and critique of the ultrasound-based assessment of the mechanical properties of the soft tissues of the foot and ankle. In light of this, the chapter also describes the methods where ultrasound was used in assessing the mechanical properties of soft tissue of the foot and ankle including the plantar fat pad, plantar fascia, and the Achilles tendon. In addition, the implications of assessing the ultrasound-based assessment of mechanical properties of soft tissue of foot and ankle in diagnosing pathological conditions such as pain, injury, tendinopathy, and diabetic foot ulceration are highlighted. Overall, the ultrasound-based measurements of the mechanical properties of the soft tissues of the foot and ankle show potential to be used for detecting the changes linked to diagnosis and prognosis of the injuries and pathological conditions such as tendinopathy, heel pain, and diabetic foot ulceration. However, considerations in terms of realistic loading suitable to the target tissue and operator dependency of the measurements need to be taken into account. In future, ultrasound-based methods of assessing the mechanical properties of the soft tissue that do not rely on making assumptions about the mechanical behavior or the mechanical characteristics of the soft tissue (e.g., linear stress-strain relationship in strain elastography or isotropic material in shear wave elastography) can be useful for more realistic assessment of the mechanical properties of the soft tissue. In addition, the use of 3D ultrasound elastography may become more popular for assessing the mechanical properties of the soft tissues of foot and ankle, allowing the assessment of tissue mechanics in all directions. Employing such methods in realistic loading conditions can facilitate more detailed analyses of the mechanical characteristics of the foot and ankle.

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[21] Ibrahim M, El Hilaly R, Taher M, Morsy A. A pilot study to assess the effectiveness of orthotic insoles on the reduction of plantar soft tissue strain. Clin Biomech 2013;28(1):68 72. [22] Cheng J-W, Tsai W-C, Yu T-Y, Huang K-Y. Reproducibility of sonographic measurement of thickness and echogenicity of the plantar fascia. J Clin Ultrasound 2012;40(1):14 19. [23] Wearing SC, Smeathers JE, Sullivan PM, Yates B, Urry SR, Dubois P. Plantar fasciitis: are pain and fascial thickness associated with arch shape and loading? Phys Ther 2007;87(8):1002 8. [24] Bygrave C, Betts R, Saxelby J. Diagnosing plantar fasciitis with ultrasound using Planscan. Foot 1998;8(3):141 6. [25] Mohseni-Bandpei MA, Nakhaee M, Mousavi ME, Shakourirad A, Safari MR, Vahab Kashani R. Application of ultrasound in the assessment of plantar fascia in patients with plantar fasciitis: a systematic review. Ultrasound Med Biol 2014;40(8):1737 54. [26] Radwan A, Wyland M, Applequist L, Bolowsky E, Klingensmith H, Virag I. Ultrasonography, an effective tool in diagnosing plantar fasciitis: a systematic review of diagnostic trials. Int J Sports Phys Ther 2016;11(5):663 71. [27] Grigg NL, Stevenson NJ, Wearing SC, Smeathers JE. Incidental walking activity is sufficient to induce time-dependent conditioning of the Achilles tendon. Gait Posture 2010;31(1):64 7. [28] Wearing SC, Grigg NL, Hooper SL, Smeathers JE. Conditioning of the Achilles tendon via ankle exercise improves correlations between sonographic measures of tendon thickness and body anthropometry. J Appl Physiol 2011;110(5):1384 9. [29] Wearing SC, Hooper SL, Grigg NL, Nolan G, Smeathers JE. Overweight and obesity alters the cumulative transverse strain in the Achilles tendon immediately following exercise. J Bodyw Mov Ther 2013;17(3):316 21. [30] Grigg NL, Wearing SC, Smeathers JE. Achilles tendinopathy has an aberrant strain response to eccentric exercise. Med Sci Sport Exerc 2012;44(1):12 17. [31] Naemi R, Chockalingam N. Mathematical models to assess foot-ground interaction: an overview. Med Sci Sports Exerc 2013;45(8):1524 33. [32] Hsu CC, Tsai WC, Shau YW, Lee KL, Hu CF. Altered energy dissipation ratio of the plantar soft tissues under the metatarsal heads in patients with type 2 diabetes mellitus: a pilot study. Clin Biomech 2007;22(1):67 73. [33] Hsu TC, Wang CL, Shau YW, Tang FT, Li KL, Chen CY. Altered heel-pad mechanical properties in patients with type 2 diabetes mellitus. Diabet Med 2000;17(12):854 9. [34] Hsu T, Lee Y, Shau Y. Biomechanics of the heel pad for type 2 diabetic patients. Clin Biomech 2002;17:291 6. [35] Klaesner JW, Hastings MK, Zou D, Lewis C, Mueller MJ. Plantar tissue stiffness in patients with diabetes mellitus and peripheral neuropathy. Arch Phys Med Rehabil 2002;83(12):1796 801. [36] Spears IR, Miller-Young JE. The effect of heel-pad thickness and loading protocol on measured heel-pad stiffness and a standardized protocol for inter-subject comparability. Clin Biomech 2006;21(2):204 12. [37] Zheng YP, Choi YK, Wong K, Chan S, Mak AF. Biomechanical assessment of plantar foot tissue in diabetic patients using an ultrasound indentation system. Ultrasound Med Biol 2000;26(3):451 6. [38] Erdemir A, Viveiros ML, Ulbrecht JS, Cavanagh PR. An inverse finite-element model of heel-pad indentation. J Biomech 2006;39 (7):1279 86. [39] Behforootan S, Chatzistergos PE, Chockalingam N, Naemi R. A clinically applicable non-invasive method to quantitatively assess the viscohyperelastic properties of human heel pad, implications for assessing the risk of mechanical trauma. J Mech Behav Biomed Mater 2017;68:287 95. [40] Behforootan S, Chatzistergos PE, Chockalingam N, Naemi R. A simulation of the viscoelastic behaviour of heel pad during weight-bearing activities of daily living. Ann Biomed Eng 2017;45(12):2750 61. [41] Chatzistergos PE, Naemi R, Sundar L, Ramachandran A, Chockalingam N. The relationship between the mechanical properties of heel-pad and common clinical measures associated with foot ulcers in patients with diabetes. J Diabetes Complicat 2014;28(4):488 93. [42] Chatzistergos PE, Naemi R, Chockalingam N. A method for subject-specific modelling and optimisation of the cushioning properties of insole materials used in diabetic footwear. Med Eng Phys 2015. [43] Hsu TC, Wang CL, Tsai WC, Kuo JK, T. FT. Comparison of the mechanical properties of the heel pad between young and elderly adults. Arch Phys Med Rehabil 1998;79:1101 4. [44] Parker D, et al. A device for characterising the mechanical properties of the plantar soft tissue of the foot. Med Eng Phys 2015;37 (11):1098 104. [45] Naemi R, Chatzistergos PE, Chockalingam N. A mathematical method for quantifying in vivo mechanical behaviour of heel pad under dynamic load. Med Biol Eng Comput 2016;54(2 3):341 50. [46] Behforootan S, Chatzistergos P, Naemi R, Chockalingam N. Finite element modelling of the foot for clinical application: a systematic review. Med Eng Phys 2017;39:1 11. [47] Ahanchian N, Nester CJ, Howard D, Ren L, Parker D. Estimating the material properties of heel pad sub-layers using inverse finite element analysis. Med Eng Phys 2017;40:11 19. [48] Zheng YP, Huang YP, Zhu YP, Wong M, He JF, Huang ZM. Development of a foot scanner for assessing the mechanical properties of plantar soft tissues under different bodyweight loading in standing. Med Eng Phys 2012;34(4):506 11. [49] Arampatzis A, Stafilidis S, DeMonte G, Karamanidis K, Morey-Klapsing G, Bru¨ggemann GP. Strain and elongation of the human gastrocnemius tendon and aponeurosis during maximal plantarflexion effort. J Biomech 2005;38(4):833 41. [50] Bohm S, Mersmann F, Marzilger R, Schroll A, Arampatzis A. Asymmetry of Achilles tendon mechanical and morphological properties between both legs. Scand J Med Sci Sport 2015;25(1):e124 32.

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[51] Chimenti RL, et al. Altered tendon characteristics and mechanical properties associated with insertional achilles tendinopathy. J Orthop Sport Phys Ther 2014;44(9):680 9. [52] Kubo K, Miyazaki D, Shimoju S, Tsunoda N. Relationship between elastic properties of tendon structures and performance in long distance runners. Eur J Appl Physiol 2015;115(8):1725 33. [53] Maganaris CN, Paul JP. Human tendon mechanical properties. J Physiol 1999;1999:307 13. [54] de Oliveira LF, Peixinho CC, Silva GA, Menegaldo LL. In vivo passive mechanical properties estimation of Achilles tendon using ultrasound. J Biomech 2016;49(4):507 13. [55] Peltonen J, Cronin NJ, Stenroth L, Finni T, Avela J. Achilles tendon stiffness is unchanged one hour after a marathon. J Exp Biol 2012;215 (20):3665 71. [56] Maganaris CN, Paul JP. Load-elongation characteristics of in vivo human tendon and aponeurosis. J Exp Biol 2000;203(Pt 4):751 6. [57] Maganaris CN. Tensile properties of in vivo human tendinous tissue. J Biomech 2002;35(8):1019 27. [58] Farris DJ, Trewartha G, McGuigan MP, Lichtwark GA. Differential strain patterns of the human Achilles tendon determined in vivo with freehand three-dimensional ultrasound imaging. J Exp Biol 2013;216(4):594 600. [59] Obst SJ, Renault J-B, Newsham-West R, Barrett RS. Three-dimensional deformation and transverse rotation of the human free Achilles tendon in vivo during isometric plantarflexion contraction. J Appl Physiol 2014;116(4):376 84. [60] Lichtwark GA, Wilson AM. In vivo mechanical properties of the human Achilles tendon during one-legged hopping. J Exp Biol 2005;208 (24):4715 25. [61] Bogaerts S, Desmet H, Slagmolen P, Peers K. Strain mapping in the Achilles tendon a systematic review. J Biomech 2016;49(9):1411 19. [62] Seynnes OR, et al. Ultrasound-based testing of tendon mechanical properties: a critical evaluation. J Appl Physiol 2015;118(2):133 41. [63] Lin C-Y, Chen P-Y, Shau Y-W, Tai H-C, Wang C-L. Spatial-dependent mechanical properties of the heel pad by shear wave elastography. J Biomech 2017;53:191 5. [64] Lin C-Y, Lin C-C, Chou Y-C, Chen P-Y, Wang C-L. Heel pad stiffness in plantar heel pain by shear wave elastography. Ultrasound Med Biol 2015;41(11):2890 8. [65] Scha¨fer G, Dobos G, Lu¨nnemann L, Blume-Peytavi U, Fischer T, Kottner J. Using ultrasound elastography to monitor human soft tissue behaviour during prolonged loading: a clinical explorative study. J Tissue Viability 2015;24(4):165 72. [66] Naemi R, Chatzistergos P, Suresh S, Sundar L, Chockalingam N, Ramachandran A. Can plantar soft tissue mechanics enhance prognosis of diabetic foot ulcer? Diabetes Res Clin Pract 2017;126:182 91. [67] Wu C-H, Chang K-V, Mio S, Chen W-S, Wang T-G. Sonoelastography of the plantar fascia. Radiology 2011;259(2):502 7. [68] Wu C-H, Chen W-S, Wang T-G. Plantar fascia softening in plantar fasciitis with normal B-mode sonography. Skelet Radiol 2015;44 (11):1603 7. [69] Lee S-Y, et al. Ultrasound elastography in the early diagnosis of plantar fasciitis. Clin Imaging 2014;38(5):715 18. [70] Kapoor A, Sandhu HS, Sandhu PS, Kapoor A, Mahajan G, Kumar A. Realtime elastography in plantar fasciitis: comparison with ultrasonography and MRI. Curr Orthop Pract 2010;21(6):600 8. [71] Sconfienza LM, et al. Real-time sonoelastography of the plantar fascia: comparison between patients with plantar fasciitis and healthy control subjects. Radiology 2013;267(1):195 200. [72] Kim M, Choi YS, You MW, Kim JS, Young KW. Sonoelastography in the evaluation of plantar fasciitis treatment. Ultrasound Q 2016;32 (4):327 32. [73] Aubry S, et al. Biomechanical properties of the calcaneal tendon in vivo assessed by transient shear wave elastography. Skelet Radiol 2013;42 (8):1143 50. [74] Siu W, Chan C, Lam C, Lee C, Ying M. Sonographic evaluation of the effect of long-term exercise on Achilles tendon stiffness using shear wave elastography. J Sci Med Sport 2016;19(11):883 7. 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Diagnostic performance of axial-strain sonoelastography in confirming clinically diagnosed achilles tendinopathy: comparison with B-mode ultrasound and color Doppler imaging. Ultrasound Med Biol 2015;41(1):15 25. [80] Klauser AS, et al. Achilles tendon assessed with sonoelastography: histologic agreement. Radiology 2013;267(3):837 42. [81] Evranos B, Idilman I, Ipek A, Polat SB, Cakir B, Ersoy R. Real-time sonoelastography and ultrasound evaluation of the Achilles tendon in patients with diabetes with or without foot ulcers: a cross sectional study. J Diabetes Complicat 2015;29(8):1124 9. [82] Zhang L, et al. Evaluation of elastic stiffness in healing achilles tendon after surgical repair of a tendon rupture using in vivo ultrasound shear wave elastography. Med Sci Monit 2016;22:1186 91. [83] Tan S, et al. Real-time sonoelastography of the Achilles tendon: pattern description in healthy subjects and patients with surgically repaired complete ruptures. 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[84] Prado-Costa R, Rebelo J, Monteiro-Barroso J, Preto AS. Ultrasound elastography: compression elastography and shear-wave elastography in the assessment of tendon injury. Insights Imaging 2018;9(5):791 814. [85] Rı´os-Dı´az J, Martı´nez-Paya´ JJ, del Ban˜o-Aledo ME, de Groot-Ferrando A, Botı´a-Castillo P, Ferna´ndez-Rodrı´guez D. Sonoelastography of Plantar fascia: reproducibility and pattern description in healthy subjects and symptomatic subjects. Ultrasound Med Biol 2015;41 (10):2605 13. [86] Chatzistergos PE, Behforootan S, Allan D, Naemi R, Chockalingam N. Shear wave elastography can assess the in-vivo nonlinear mechanical behavior of heel-pad. J Biomech 2018;.

Chapter 21

3D Surface Scanning of the Foot and Ankle Scott Telfer1,2,3 1

Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States, 2Department of Mechanical Engineering,

University of Washington, Seattle, WA, United States, 3RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States

Abstract The external shape of the foot and ankle can provide indications about its underlying health and function, and it is important when designing orthotic interventions for pathologies affecting the lower limb. Until recently, assessing the overall shape of the foot has been a largely subjective task, with only a few linear or circumferential measurements used, or time-consuming and unreliable plaster casts required to capture the full shape of the foot. 3D surface scanning allows a detailed digital model of the whole foot and ankle to be easily captured, and this model can subsequently be utilized and analyzed in numerous ways. Many technologies fall under the umbrella of 3D surface scanning, and several have been shown to be fast, accurate, and reproducible in the context of foot and ankle research. Given this, 3D surface scanning is being increasingly utilized in studies involving custom footwear and orthotics. The technology lends itself to capturing large sets of data and has been employed in several epidemiological studies. This chapter covers the development of 3D surface scanning, its foot and ankle biomechanics-related applications, and potential future uses.

21.1

Introduction

The external shape of the foot and ankle can provide significant information about its function and health. Indeed, most clinical examinations will begin with a simple visual exam that is intended to identify any obvious problems, such as deformity or swelling [1]. Three-dimensional (3D) surface scanning, sometimes also known as reality capture, covers a set of technologies that allow the shape of an object to be digitally recorded and used for further applications. At the time of writing, 3D scanning has found several clinical applications in the healthcare field, including craniomaxillofacial surgery [2,3], dentistry [4], and prosthetics/orthotics [5]. In the context of foot and ankle health, 3D scanning has several advantages over a purely visual inspection, among others: scans are fast to capture, they can be analyzed in detail and compared at different time points, changes can be objectively measured, and the digital files can be stored and revisited. A foot scan (Fig. 21.1) can also be integrated into a digital manufacturing workflow and be used to inform the design of custom medical devices, such as orthotic interventions. Beyond healthcare applications, the technology has also been utilized for ergonomic factors in footwear design [6]. A 2010 review of 3D surface scanning of the human foot by the author stated that “Modern 3D surface scanning systems can obtain accurate and repeatable digital representations of the foot shape and have been successfully used in medical, ergonomic, and footwear development applications” [7]. Further development of the technology, in terms of reductions in hardware cost, ease of use, and integration with downstream applications have occurred in the years since this publication. This is in addition to a significant number of additional peer-reviewed publications related to foot and ankle biomechanics that have utilized this technology in some form (see detailed discussion below). This chapter aims to: (1) provide an overview of the development of 3D surface scanning technologies; (2) review how it has been applied in the context of foot and ankle biomechanics; and (3) propose areas of future research where it may be further utilized. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00019-6 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 21.1 Foot scan with color texture. Note artifacts in the scan between the toes, likely a result of lack of resolution in this region.

FIGURE 21.2 Volume of results for “3D Scanning” and related search terms in books sampled by the Google Ngram Viewer Tool between 1970 and 2019.

21.1.1 The history and development of 3D scanning The reduction in cost of consumer level computers and electronics has, in the last decade or so, led to an increase in the general awareness of 3D surface scanning (Fig. 21.2). Prior to and beyond its applications in the healthcare arena, the

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technique has found roles in several diverse industries, from production and civil engineering [8] to fashion and art [9]. Initial work in the area of scanning body parts was driven in part by the entertainment industry, particularly for animation [10], with one of the earliest systems being a head scanner developed for this market by Cyberware Laboratories (Los Angeles, CA).

21.1.2 Technologies In general, a 3D scanner will collect a point cloud of data, with each point representing a discrete location on the surface of the object. Depending on the capabilities of the scanner, the object color for each point may also be captured at this time (Fig. 21.3). To make the dataset more useful, for example, allowing it to be imported into computer aided design (CAD) software later in the workflow, this point cloud may be reconstructed into a surface model, usually in the form of a polygon mesh (.stl and.obj formats are most common). At the time of writing, the majority of commercially available 3D scanners now include this functionality within their data capture software. The density and accuracy of the point cloud depends on several factors relating to the technology used, and indeed one of the early limiting factors for noncontact 3D surface scanning was being able to store the relatively large data files, a problem that has since been overcome by advances in computer hardware. Processing the scan data to remove noise is usually necessary, especially for optical systems, and this can cause some detail to be lost due to smoothing. Systems range dramatically in cost, with simple, handheld consumer-level systems being available for a few hundred dollars, to larger, dedicated units designed to capture weightbearing scans of the full foot and ankle costing tens of thousands of dollars also available. Several scans may be taken of the same object from different viewpoints. These can then be automatically or manually registered using overlapping areas and then joined together to create a full model. Handheld scanners collect data in relation to an internal coordinate system, therefore when the system is in motion its position must be determined. This can be achieved using reference features on the object being scanned, or by an external tracking method such as magnetic sensors. There are a range of techniques that can and have been used to obtain a digital 3D model of the shape of the foot and ankle. These are too numerous to comprehensively review in this space, but the key methods are briefly covered. Primarily, systems that can be used to scan the foot (or foot and ankle) directly will be discussed; however, it should be noted that there are perhaps an even greater number of desktop scanning systems available that can be used to scan foam box or plaster cast impressions of the foot [11]. In practice, 3D scanners can be divided into contact or noncontact scanners [12]:

21.1.2.1 Contact scanners (coordinate measuring machines) Contact scanners use single or multiple probes, usually physically connected via articulating link segments, that are placed against the object to be scanned. The Orthema CAD/CAM system (Orthofit Verkaufs GmbH; Rotkreuz, Switzerland) is an example of this, which used a set of probes to touch the plantar surface of the foot, and depth measurements taken from each. A limitation of this type of system is that the contact between the probes and the foot may deform the soft tissues, although depending on the application this error may be acceptable. Generally, these systems have been surpassed by their noncontact counterparts that have the advantage of being able to capture millions of data points in a very short time (usually a few seconds), several orders of magnitude greater than the datasets that can reasonably be captured by a contact scanner. Until recently, contact scanners had the advantage of being more accurate, but improvements in noncontact technologies have largely overcome this in recent years, at least to the levels required for most foot and ankle measurements [13].

21.1.2.2 Noncontact scanners Similar to a camera, a noncontact 3D scanner collects information about an object within its field of view; however, the key information relates to the distance the points are from the device. Noncontact scanners have the advantage of FIGURE 21.3 Scan of talus. (A) point cloud; (B) wireframe mesh; (C) color texture applied.

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avoiding potential errors from tissue distortion associated with physical probes or other manual measurement tools. These types of scanners can be further subclassified as active or passive, where some form of light is projected onto the object (active) or the natural lighting of the environment is used (passive). Time-of-flight systems generally use a laser light to probe the object. The systems use a laser rangefinder to determine the time it takes for a pulse of light to reach the object and reflect back to the scanner. The accuracy of these systems is determined by how precisely the time-of-flight can be measured (for context: it takes approximately 3 picoseconds [10 12] for a pulse of light to travel 1 mm). Hardware requirements to accurately measure these times limit the use of these systems for close range scanning. Because of this, the technology is rarely used for foot scanning applications. Triangulation systems, such as the FastSCAN (Polhemus, Colchester, Vermont) use a laser dot or line to probe the object of interest. In this case, a camera, mounted at a fixed and known position near the laser emitter, is used to find the location of the dot or line. The distance the laser strikes the object at can be calculated from its position in the camera’s field of view. Structured light scanners, such as the Solescan Blanka (STT Systems, San Sebastian, Spain), project a light pattern in to the object of interest. A camera at a known position from the projector captures the image of the light pattern and the deformation of the pattern is used to determine the shape of the object. The main advantage of this approach is speed, in that the entire field of view can be captured in a single image. Depth cameras are similar to structured light scanners, but these use an infrared emitter to project a pattern onto the object, and have an infrared camera to capture its distortion. This technology is used in the handheld Sense 3D scanner (3D Systems, Rock Hill, SC), and was used in the now discontinued Kinect sensor (Microsoft Corporation., Redmond, WA). Photogrammetry is a passive technique used by most smartphone or tablet 3D scanning apps and uses images of the object of interest taken from several different angles. By looking for features that appear in multiple images, processing software is used to determine the position of the camera when the image was taken, and subsequently reconstruct a 3D model. Fine details are often lost when using this technique, and the process can be computationally expensive. However, it does not require any specialist equipment, only a camera, and results can be acceptable [14]. Dynamic 3D scanning. One of the latest developments in 3D scanning is the ability to capture the shape of the foot dynamically. As a piece of anatomy that is in most cases flexible and deforms considerably during weightbearing, one of the limitations of using a digital model of the foot for many of the applications described in this chapter is that only a single, static digital model is used. By capturing the shape of the foot dynamically, using one or a combination of the techniques described above, some of these limitations can be overcome. This approach has been used in the context of markerless motion capture, circumventing the need for markers to be attached to the skin of the foot for multisegment foot models [15]. Other investigators have used dynamic 3D scanning to study the plantar surface of the foot during gait, producing distance maps they considered to be analogs to plantar pressure distribution measurements [16]. The repeatability of one such system has been shown to be around 3 mm between days for dynamic trials [17].

21.2

Foot-specific applications and considerations

This section will cover research performed using 3D scanners in areas related to foot and ankle biomechanics, including reproducibility studies, population studies, footwear and orthotic studies, musculoskeletal models, and internal anatomy.

21.2.1 Reliability and comparisons to other techniques The question of what an acceptable level of accuracy in the context of a scan of the foot and ankle likely depends on the specific application. Many scanning system manufacturers stated accuracies are, in the author’s experience, only achievable under highly controlled conditions. In particular, capturing sharp edges, for example toenails or edges of bones, can be difficult, and several systems are affected by the lighting conditions in the scanning environment. Several researchers have investigated the reliability of different foot and ankle scanning systems (Table 21.1). Linear, angular, and circumferential measurements may be the primary concern, however the reliability of the overall shape capture is also of interest, and each require different approaches. In general, 3D surface scanning is reliable and valid, whether in the context of general measurements of the foot and ankle, or for the design of orthotics.

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TABLE 21.1 Studies investigating the reliability of 3D surface scanning of the foot and ankle. Author (year)

System(s)/techniques

Measurements

Findings

Witana [18]

YETI I foot scanner (Vorum Research Corporation, Vancouver, BC); laser triangulation system

Selected linear and circumferential measurements

With the exception of heel width, there were no significant differences between the measurements determined manually vs those determined from the scan data

Carroll et al. [19]

Virtual Orthotics noncontact scanner (Virtual Orthotics, NSW, Australia); structured light system

Selected linear measurements and forefoot to hindfoot alignment

3D scanning is reliable, independent of clinical experience, and variability was reduced compared to neutral suspension casting

De Mits [20]

INFOOT USB (standard type; IWare Laboratory Co Ltd, Japan); laser triangulation system

Repeated scans of the same set of feet and comparison to clinical measurements

Good validity and reliability was shown for most of the measurements in patients with rheumatoid arthritis

Telfer [11]

Easy Foot Scan (Baltic Orthoservice, UAB, Kaunas, Lithuania) Direct 3D scans (full and partial weightbearing); laser triangulation system

Selected linear measurements and volume overlap

Foot orthotics designed using 3D scanned foot models were equivalent to those obtained using traditional plaster casting methods

Lee [13]

INFOOT USB, (IFU-S-01, I-Ware Laboratory Co., Ltd, Japan); laser triangulation system

Foot length Ball of foot length Outside ball of foot length Foot breadth diagonal Foot breadth horizontal Heel breadth

The majority of measurements obtained using the 3D scanner were more accurate and precise than those obtained using conventional methods

21.2.2 Population studies Most modern foot scanning systems allow a digital model of the foot to be captured in just a few seconds. This feature lends the technology to large scale population studies. This can allow variation across different groups of individuals, driven perhaps by genetic or environmental factors, to be examined. These types of population studies are useful, for example, in identifying varying ergonomic parameters for footwear design, and can be used by manufacturers to tailor their size ranges to different populations. In such a study, Lee et al. [21] found differences between the shapes of Taiwanese and Japanese feet [21]. 3D scanning has also been used to study differences in foot shape between populations of habitually shod and unshod runners [22]. The authors compared a group of 168 unshod Indian runners to 196 shod Chinese runners and found differences in several measurements, including hallux angle for the shod group. While this study design does not allow a casual effect on hallux angle, it does add to an increasing body of evidence suggesting the footwear may play a role in the development of hallux valgus. Scan data have been used to assess associations between demographic factors and differences in foot shape variables. The effects of sex have been tested in several studies [23 25]. In a sample of 291 older Japanese adults it was found that, after normalization to foot length, instep and navicular height was greater in males, and the transverse plane angle of the hallux was greater in females [23]. Similar results for the influence of sex were found in a study of Europeans, with this study further subcategorizing their study population into foot types described as voluminous, flatpointed, and slender [24]. Similarly, the effect of body mass index on foot shape has been assessed [26]. Obese participants were found to have foot shapes that were significantly wider than their healthy and overweight counterparts. In addition, the authors noted that the changes in foot shape were not always reflected by measurements using standard anatomical landmarks, and that more comprehensive foot shape data such as that generated by 3D scanning systems was necessary to inform footwear design. Lee and Wang [27] studied 3000 Taiwanese individuals and were able to identify 3 foot shape types for both male and female feet using surface scan data [27]. The effect of age has also been studied, the foot shape of older Romanian female adults being found to be significantly different from the general

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female population in that country [28]. Surface scanning can also be used to study differences in foot and ankle shape between weightbearing and nonweightbearing conditions [29 31].

21.2.3 Orthoses and footwear Arguably the area where 3D scanning of the foot and ankle has been most fully utilized is in the design of custom or customized footwear and orthotic devices. Traditional, artisan techniques for capturing the foot shape and subsequently producing a modified positive cast suitable for constructing the device against tend to be messy (usually requiring dedicated plaster rooms where negatives and positives are made and modified) and lack the repeatability offered by a fully digital process. The scanning process is clean and the subsequent results can be stored and easily communicated electronically. Digital scans also allow integration with other emerging manufacturing techniques such as 3D printing, which presents several opportunities for innovative device development [32]. Overall, the digital process has been shown to cost considerably less than traditional, manual methods of obtaining the shape of the foot [33]. Several studies have assessed shoe fit using measurements from 3D scanners, particularly for older people. By comparing dimensions derived from the foot scans to the dimensions of the shoe lasts, it was shown that shoes could be appropriately chosen for individuals using standard, linear measurement techniques [6]. The fit of protective footwear for work has also been assessed for older adults [34]. In this study, several linear measurements were taken from 3D surface scans and compared to the corresponding measurements in protective footwear. The authors found that this was an effective method that may help with improving the design of industrial footwear. Mickle et al. studied the foot morphology of over 300 older adults and concluded that footwear for older people should be designed with the altered morphology of this population in mind [35]. Several studies of insoles and therapeutic footwear have utilized 3D scanning. This has included studies of patient populations with diabetic foot disease [36 38] or rheumatoid arthritis [39], with both direct scans and scans of foam or plater impressions being used. Witana et al. used 3D scanning to quantify the deformation of the foot while participants stood on support surfaces with different cushioning properties, and found that the midfoot structure could change shape, independent of the hindfoot and forefoot regions [40]. The effect of heel height has also been studied using 3D scanning, with scans taken with 5 and 10 cm heel wedges found to elevate the fourth and fifth metatarsophalangeal joints and cause the small toe to move laterally compared to foot flat [41]. Beyond insoles and footwear, the use of 3D scanning to produce custom ankle foot orthosis (AFO) designs has also been explored [42]. These require a larger area of anatomy to be captured, needing the foot and leg rather than just the foot itself, meaning that structured light handheld scanners are most commonly used for this application. This has been particularly useful in studies proposing new designs of AFOs [32,43]. A randomized controlled trial has been performed, comparing AFOs produced using laser scanning to those produced using traditional casting methods [44]. No significant differences in the time were seen, and more fitting problems were reported with the AFOs based on scans than those based on casts, reflecting the additional complexity involved in prescribing these devices. The authors did note that significant technological improvements in scanning and fabrication had been achieved since they initiated the trial, and likely further testing is required to take these advances into account. There are limitations with designing footwear and orthoses from scans that should be noted. Polygon mesh formats can result in large files that are not particularly easy to edit, therefore the use of digital models of the foot for footwear and orthosis design has required the development of software that can imitate the modifications made by the orthotist on a physical model of the foot, along with the operator skills to effectively perform these modifications (Fig. 21.4).

21.2.4 Musculoskeletal models To make musculoskeletal models of the foot more easily accessible, the Anybody modeling platform includes an option to use a 3D scan of the foot to scale their foot model [45]. This is achieved by having the user identify anatomical landmarks on the digital model. This approach is likely less accurate than having a CT or other medical imaging scan of the foot, or even manual palpation, but is much easier to obtain.

21.2.5 Bones and other internal anatomy Three-dimensional scanning technologies have been used in studies of internal bony anatomy taken from cadaver specimens [46], and to quantify visible areas of bone for different surgical approaches [47]. Determining mean and variations

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FIGURE 21.4 Generating a foot orthosis from a scanned foot impression. (A) foam box impression; (B) scanned image of box; (C) orthosis being designed in CAD program; (D) final design.

in surface anatomy for the design of implants is a useful example of this [48]. These scans can also be 3D printed and used as educational models [49].

21.3

Areas of future biomechanical research

As described earlier, 3D scanning of the foot and ankle has been or is at least well on its way to successful translation for several clinical applications. There is, however, considerable potential for its use to be expanded further, both in the research and clinical domains. For example, dynamic 3D scanning has been shown to be technically feasible, but a true clinical application has still to be demonstrated. By capturing the changing shape of the foot in its functional state it may be possible to better customize the footwear for comfort and biomechanical effects. Such a workflow has yet to be fully demonstrated however, and may require the scan data to be combined with other technologies to account for the soft tissues effects, however this is certainly worth exploring. The relationship between shape-based measurements and foot function has not been explored. The underlying structure of the foot has been found to be only partially associated with functional measures, however with larger studies, it may be possible to gain further insights. Tracking changes over time for injuries, for example, monitoring regional swelling after ankle strains is another area that has not been explored, and could potentially provide more information than currently used measurement. The ability to capture foot shape using a smart phone app opens a host of opportunities for large-scale studies on the general population. This could, for example, allow the prevalence of deformities that are easily detectable from foot shape, such as hallux valgus, to be determined in a larger sample than has previously been possible. 3D scanners can also capture color, and this has not been explored. Changes in color can indicate blood supply, and this may be a potential area of future investigations.

References [1] Alazzawi S, Sukeik M, King D, Vemulapalli K. Foot and ankle history and clinical examination: a guide to everyday practice. World J Orthop 2017;8:21 9. Available from: https://doi.org/10.5312/wjo.v8.i1.21. [2] Knoops PGM, Beaumont CAA, Borghi A, Rodriguez-Florez N, Breakey RWF, Rodgers W, et al. Comparison of three-dimensional scanner systems for craniomaxillofacial imaging. J Plast Reconstr Aesthetic Surg 2017;70:441 9. Available from: https://doi.org/10.1016/j. bjps.2016.12.015.

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[3] Kuijpers MAR, Chiu Y-T, Nada RM, Carels CEL, Fudalej PS. Three-dimensional imaging methods for quantitative analysis of facial soft tissues and skeletal morphology in patients with orofacial clefts: a systematic review. PLoS One 2014;9:e93442. Available from: https://doi.org/ 10.1371/journal.pone.0093442. [4] Ireland AJ, McNamara C, Clover MJ, House K, Wenger N, Barbour ME, et al. 3D surface imaging in dentistry what we are looking at. Br Dent J 2008;205:387 92. Available from: https://doi.org/10.1038/sj.bdj.2008.845. [5] Geil MD. Consistency, precision, and accuracy of optical and electromagnetic shape-capturing systems for digital measurement of residual-limb anthropometrics of persons with transtibial amputation. J Rehabil Res Dev 2007;44:515 24. [6] Menz HB, Auhl M, Ristevski S, Frescos N, Munteanu SE. Evaluation of the accuracy of shoe fitting in older people using three-dimensional foot scanning. J Foot Ankle Res 2014;7:3. 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[11] Telfer S, Gibson KS, Hennessy K, Steultjens MP, Woodburn J. Computer-aided design of customized foot orthoses: reproducibility and effect of method used to obtain foot shape. Arch Phys Med Rehabil 2012;93:863 70. Available from: https://doi.org/10.1016/j.apmr.2011.12.019. [12] Curless B. Brian, From range scans to 3D models. ACM SIGGRAPH Comput Graph 1999;33:38 41. Available from: https://doi.org/10.1145/ 345370.345399. [13] Lee Y-C, Lin G, Wang M-JJ. Comparing 3D foot scanning with conventional measurement methods. J Foot Ankle Res 2014;7:44. Available from: https://doi.org/10.1186/s13047-014-0044-7. [14] Petriceks AH, Peterson AS, Angeles M, Brown WP, Srivastava S. Photogrammetry of human specimens: an innovation in anatomy education. J Med Educ Curric Dev 2018;5. Available from: https://doi.org/10.1177/2382120518799356 238212051879935. [15] Van den Herrewegen I, Cuppens K, Broeckx M, Barisch-Fritz B, Vander Sloten J, Leardini A, et al. Dynamic 3D scanning as a markerless method to calculate multi-segment foot kinematics during stance phase: methodology and first application. J Biomech 2014;47:2531 9. Available from: https://doi.org/10.1016/j.jbiomech.2014.06.010. [16] Samson W, Van Hamme A, Sanchez S, Che`ze L, Van Sint Jan S, Feipel V. Foot roll-over evaluation based on 3D dynamic foot scan. Gait Posture 2014;39:577 82. Available from: https://doi.org/10.1016/j.gaitpost.2013.09.014. [17] Thabet AK, Trucco E, Salvi J, Wang W, Abboud RJ. Dynamic 3D shape of the plantar surface of the foot using coded structured light: a technical report. J Foot Ankle Res 2014;7:5. Available from: https://doi.org/10.1186/1757-1146-7-5. [18] Witana CP, Xiong S, Zhao J, Goonetilleke RS. Foot measurements from three-dimensional scans: a comparison and evaluation of different methods. Int J Ind Erg 2006;36:789 807. Available from: https://doi.org/10.1016/j.ergon.2006.06.004. [19] Carroll M, Annabell M-E, Rome K. Reliability of capturing foot parameters using digital scanning and the neutral suspension casting technique. J Foot Ankle Res 2011;4:9. Available from: https://doi.org/10.1186/1757-1146-4-9. [20] De Mits S, Coorevits P, De Clercq D, Elewaut D, Woodburn J, Roosen P. Reliability and validity of the INFOOT three-dimensional foot digitizer for patients with rheumatoid arthritis. J Am Podiatr Med Assoc 2011;101:198 207. [21] Lee Y-C, Kouchi M, Mochimaru M, Wang M-J. Comparing 3d foot shape models between taiwanese and japanese females. J Hum Ergol (Tokyo) 2015;44:11 20. [22] Shu Y, Mei Q, Fernandez J, Li Z, Feng N, Gu Y. Foot morphological difference between habitually shod and unshod runners. PLoS One 2015;10:e0131385. Available from: https://doi.org/10.1371/journal.pone.0131385. [23] Saghazadeh M, Kitano N, Okura T. Gender differences of foot characteristics in older Japanese adults using a 3D foot scanner. J Foot Ankle Res 2015;8:29. Available from: https://doi.org/10.1186/s13047-015-0087-4. [24] Krauss I, Grau S, Mauch M, Maiwald C, Horstmann T. Sex-related differences in foot shape. Ergonomics 2008;51:1693 709. Available from: https://doi.org/10.1080/00140130802376026. [25] Luo G, Houston VL, Mussman M, Garbarini M, Beattie AC, Thongpop C. Comparison of male and female foot shape. J. Am. Podiatr. Med. Assoc. n.d.;99:383 90. [26] Price C, Nester C. Foot dimensions and morphology in healthy weight, overweight and obese males. Clin Biomech 2016;37:125 30. Available from: https://doi.org/10.1016/j.clinbiomech.2016.07.003. [27] Lee Y-C, Wang M-J. Taiwanese adult foot shape classification using 3D scanning data. Ergonomics 2015;58:513 23. Available from: https:// doi.org/10.1080/00140139.2014.974683. [28] Sarghie B, Mihai A, Costea M, Rezu¸s E. Comparative anthropometric study regarding the foot of elderly female population. Procedia Eng 2017;181:182 6. Available from: https://doi.org/10.1016/J.PROENG.2017.02.367. [29] Houston VL, Luo G, Mason CP, Mussman M, Garbarini M, Beattie AC. Changes in male foot shape and size with weightbearing. J. Am. Podiatr. Med. Assoc. n.d.;96:330 43. [30] Xiong S, Goonetilleke RS, Zhao J, Li W, Witana CP. Foot deformations under different load-bearing conditions and their relationships to stature and body weight. Anthropol Sci 2009;117:77 88. Available from: https://doi.org/10.1537/ase.070915.

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[31] Tsung BYS, Zhang M, Fan YB, Boone DA. Quantitative comparison of plantar foot shapes under different weight-bearing conditions. J. Rehabil. Res. Dev. n.d.;40:517 26. [32] Telfer S, Pallari J, Munguia J, Dalgarno K, McGeough M, Woodburn J. Embracing additive manufacture: implications for foot and ankle orthosis design. BMC Musculoskelet Disord 2012;13:84. Available from: https://doi.org/10.1186/1471-2474-13-84. [33] Payne C. Cost Benefit comparison of plaster casts and optical scans of the foot for the manufacture of foot orthoses. Aust J Pod Med 2007;41:29 31. [34] Irzma´nska E, Okrasa M. Evaluation of protective footwear fit for older workers (60 1 ): a case study using 3D scanning technique. Int J Ind Erg 2018;67:27 31. Available from: https://doi.org/10.1016/J.ERGON.2018.04.001. [35] Mickle KJ, Munro BJ, Lord SR, Menz HB, Steele JR. Foot shape of older people: implications for shoe design. Footwear Sci 2010;2:131 9. Available from: https://doi.org/10.1080/19424280.2010.487053. [36] Ulbrecht JS, Hurley T, Mauger DT, Cavanagh PR. Prevention of recurrent foot ulcers with plantar pressure-based in-shoe orthoses: the CareFUL prevention multicenter randomized controlled trial. Diabetes Care 2014;37:1982 9. Available from: https://doi.org/10.2337/dc13-2956. [37] Telfer S, Woodburn J, Collier A, Cavanagh PR. Virtually optimized insoles for offloading the diabetic foot: a randomized crossover study. J Biomech 2017;60:157 61. Available from: https://doi.org/10.1016/j.jbiomech.2017.06.028. [38] Khodaei B, Saeedi H, Jalali M, Farzadi M, Norouzi E. Comparison of plantar pressure distribution in CAD CAM and prefabricated foot orthoses in patients with flexible flatfeet. Foot 2017;33:76 80. Available from: https://doi.org/10.1016/j.foot.2017.07.002. [39] Gibson KS, Woodburn J, Porter D, Telfer S. Functionally optimised orthoses for early rheumatoid arthritis foot disease: a study of mechanisms and patient experience. Arthritis Care Res (Hoboken) 2014;66:1456 64. Available from: https://doi.org/10.1002/acr.22060. [40] Witana CP, Goonetilleke RS, Xiong S, Au EYL. Effects of surface characteristics on the plantar shape of feet and subjects’ perceived sensations. Appl Erg 2009;40:267 79. Available from: https://doi.org/10.1016/j.apergo.2008.04.014. [41] Wan FKW, Yick K-L, Yu WWM. Validation of a 3D foot scanning system for evaluation of forefoot shape with elevated heels. Measurement 2017;99:134 44. Available from: https://doi.org/10.1016/J.MEASUREMENT.2016.12.005. [42] Cha YH, Lee KH, Ryu HJ, Joo IW, Seo A, Kim D-H, et al. Ankle-foot orthosis made by 3D printing technique and automated design software. Appl Bionics Biomech 2017;2017:9610468. Available from: https://doi.org/10.1155/2017/9610468. [43] Walbran M, Turner K, McDaid AJ. Customized 3D printed ankle-foot orthosis with adaptable carbon fibre composite spring joint. Cogent Eng 2016;3. Available from: https://doi.org/10.1080/23311916.2016.1227022. [44] Roberts A, Wales J, Smith H, Sampson CJ, Jones P, James M. A randomised controlled trial of laser scanning and casting for the construction of ankle foot orthoses. Prosthet Orthot Int 2016;40:253 61. Available from: https://doi.org/10.1177/0309364614550263. [45] Al-Munajjed AA, Bischoff JE, Dharia MA, Telfer S, Woodburn J, Carbes S. Metatarsal loading during gait-a musculoskeletal analysis. J Biomech Eng 2016;138:034503 034503-6. Available from: https://doi.org/10.1115/1.4032413. [46] Harper CM, Ruff CB, Sylvester AD. Gorilla calcaneal morphological variation and ecological divergence. Am. J Phys Anthropol 2020. Available from: https://doi.org/10.1002/ajpa.24135. [47] Magnusson EA, Telfer S, Jackson M, Githens MF. Does a medial malleolar osteotomy or posteromedial approach provide greater surgical visualization for the treatment of talar body fractures? J Bone Joint Surg Am 2021;103(24):2324 30. Available from: https://doi.org/10.2106/ JBJS.21.00299. [48] Kumar A, Donley B, Cavanagh PR. Design of an implant for first metatarsophalangeal hemi-arthroplasty. Comput Methods Biomech Biomed Eng 2014;17:1777 84. Available from: https://doi.org/10.1080/10255842.2013.766723. [49] Thomas DB, Hiscox JD, Dixon BJ, Potgieter J. 3D scanning and printing skeletal tissues for anatomy education. J Anat 2016;229:473 81. Available from: https://doi.org/10.1111/joa.12484.

Chapter 22

Cadaveric Gait Simulation Patrick Aubin1,2,3 and William R. Ledoux1,2,3 1

RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 2Department of

Mechanical Engineering, University of Washington, Seattle, WA, United States, 3Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Abstract Cadaveric gait simulation has evolved from static tests representing small instances in the gait cycle to fully dynamic simulations from heel strike to toe off. In this chapter, we review dynamic simulations of stance phase. We describe the techniques used to generate the simulations, specifically what inputs and outputs are employed, and what hardware is implemented to move the foot and/or the “ground.” Strategies for applying and controlling muscle forces are discussed, as are the many specialized model outputs. We also detail the various limitations of these simulators, which often requires lower loads and slower speeds than physiologic. The number of actuated muscles is often reduced, as is the number of available degrees of freedom of motion. Finally, it can be difficult to model disease states accurately. Conversely, researchers have used dynamic cadaveric models to explore a wide variety of clinical applications, including developing other hardware, tracking bone motion in various conditions, exploring muscle and ligament function, and studying joint replacements as well as other orthopedic conditions. We conclude the chapter with a discussion on future biomechanical research.

22.1

Introduction

When studying a part of the human musculoskeletal system such as the foot and ankle, researchers often undertake a multi-faceted approach. If a parameter of interest (e.g., bone kinematics) can be measured safely, accurately, and ethically on living subjects, then that should be the method of choice. However, certain experiments (e.g., the effect of malalignment of a surgical procedure) cannot be efficaciously conducted on living subjects; hence, cadaveric studies are often undertaken. It is acknowledged that cadaveric models are lacking in certain aspects (e.g., cadaveric surgeries can typically only be done once) that invite other techniques such as computational modeling, that in turn have limitations (e.g., lack of anatomical detail) that can suggest the need to consider biofidelic 3D-printed anatomical test beds. Each of these modeling approaches have their own strengths and weaknesses, and each can address an aspect of the greater biomechanical question posed. Concerning the foot and ankle, there have been many studies (too numerous to mention) that have included living subjects or cadaveric specimens or computational models; more recently, 3D printed constructs that include anatomical details have been explored [1]. Elsewhere in this book, we have devoted chapters to the aforementioned living subject studies and computational modeling of the foot and ankle. In this chapter, we review cadaveric simulations of the foot and ankle in detail. Specifically, we are interested in dynamic cadaveric gait simulation. There is a large body of literature that employs static or quasi-static cadaveric techniques to quantify foot function, but beginning with the work of Neil Sharkey’s team in the 1990s [2], a subset of research groups have developed various techniques of dynamic gait simulation. A validated dynamic test platform that can simulate the kinetics and kinematics of gait is much more useful than a static model that merely provides a snapshot of the gait cycle. That is not to impugn the many important studies of foot and ankle biomechanics that were conducted statically on cadavers; these bodies of work have addressed and continue to address, many important and clinically relevant questions. However, the purpose of this chapter is to review and discuss the dynamic gait simulators (DGSs) that have been developed over Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00033-0 © 2023 Elsevier Inc. All rights reserved.

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the last 30 years. First, we will highlight the more important gait simulators by summarizing how they function (e.g., inputs, outputs, validations, etc.). We will also review the limitations of dynamic gait simulation, before exploring the many clinical questions that have been tested with these simulators. We will end the chapter with a forward looking discussion of the future biomechanical considerations that remain to be addressed related to dynamic cadaveric gait simulations.

22.2

Techniques for dynamic gait simulation

Dynamic gait simulation is a technique to model gait using a cadaver lower limb specimen. As with all scientific models, gait simulators have model inputs (variables that are prescribed) and outputs (outcomes that are predicted by the model). The utility of DGSs lies in their ability to accurately predict model outputs for a given set of model inputs and thereby answer hypothesis-driven research questions. When designing a DGS, the researcher must decide what they want the model to predict (model outputs) and what inputs will be used to drive the model. Typical inputs are the cadaveric specimen, tibia 6-degree-of-freedom kinematics, the ground reaction force, and the extrinsic foot and ankle tendon forces. Model outputs are more variable and depend on the scientific question to be addressed. When specifying model inputs and outputs, one must be careful not to fall into a circular framework whereby an input is also thought of as an output. What makes DGSs both interesting and challenging to develop is their dynamic nature, whereby the foot and ankle are animated with tibia and/or ground motions, tibia and/or ground forces, and tendon forces. Controlling these forces and motions that act upon a cadaveric specimen is a difficult robotic task. During in vivo gait, angular displacement at the hip, knee, and ankle joints not only cause sagittal, coronal, and transverse plane rotation of the thigh, leg, and foot, but also superior/inferior, medial/lateral, and anterior/posterior translation of the joint centers through 3D space. Thus, the motion of the tibia with respect to the ground during walking has 6-degrees-of-freedom (DOF) (three rotations and three translations). DGSs typically employ electromechanical actuators to rotate and/or translate the cadaveric tibia through space in a manner that replicates in vivo tibial kinematics; alternatively, the same relative motion can be generated by precisely moving the ground relative to a fixed tibia. To precisely control the 6-DOF tibia motion, DGSs have employed (1) linear drives with a sliding carriage [2 19], (2) 6-DOF Stewart-type robotic platform [18 35], (3) linear cart type device [36,37], (4) customized robotic mechanisms [35,38 41], and (5) other techniques (Table 22.1, Figs. 22.1 22.9). During a simulation, dedicated computer hardware and software systems utilize position feedback control methods to achieve the desired motion [19,21 35]. Dynamic tendon force is created during a simulation by attaching the extrinsic tendons of the foot and ankle to computer controlled electrical [2 12,18,19,21 35,37,41], hydraulic [38 40], or pneumatic actuators [13 17,41,43]. Specialized mechanical clamps have been developed to attach the tendons to the actuators [23]. When significant tendon force is needed, such as for the Achilles tendon, freeze clamps that utilize liquid nitrogen provide exceptional tendon holding strength [49]. Depending on the nature of the experiment, tendons actuation can operate under force control, or position control. Force control tendon actuation means the actuator tracks a target tendon force [2 4,6 12,19,21 34,41,43] while position control means the proximal end of the tendon is moved through a target linear displacement [5,7]. Force controlled tendon actuators are programed to track an estimate of in vivo tendon force. Although direct measurement of tendon force in vivo has been achieved for the Achilles tendon [50], the difficulty in obtaining these direct measures means in vivo tendon force is often estimated using an electromyography to force model [2,6 9,11,12,21 33,43]. DGSs have had a variety of different model outputs associated with the wide range of outcomes that have been studied. For example, Milgrom et al. [10] and Meardon et al. [37] measured bone strain while Erdemir et al. determined the effect of skin movement artifact on fiberoptic measurement of tendon force [3]. Lee et al. [22] and Suckel et al. [40] measured joint pressure. Salb et al. investigated the instantaneous axis of rotation of the foot and ankle [51], while Ledoux et al. measured plantar pressure [23,26,28], among others. Optical motion capture camera systems [4,8,9,13,14,19,24,34,36,52], or magnetic [43] tracking systems have been employed to measure bone kinematics via bone mounted fiducials. Force plates have been employed to measure vertical and shear ground reaction forces [2,18,21,22,25,36,41,43], while pressure mats have been use to record plantar pressures [23,26,28] (Table 22.1). One of the key assets of DGSs is their ability to model different foot and ankle diseases and conditions. Over the past 30 years, groups have created models of flatfeet with posterior tibial tendon dysfunction [43], total ankle arthroplasty [30], ankle arthrodesis [31], first metatarsophalangeal arthrodesis [23], and metatarsalgia [28], among others. Some conditions are simulated through alteration of the foot and ankle bones, ligaments, or tendons, while replicating other conditions is achieved by altering the model inputs (e.g., tendon force). For example, Watanabe et al. created a

TABLE 22.1 Performance characteristics of dynamic cadaveric gait simulators. Institution

Inputs

# Tendons

DOFn

Stance (sec)

%BW

Outputs

VAPSHCSa

Tibia kinematics, VGRFk

9

6

2.7

75

Foot bone motion, plantar pressure

Tibia kinematics, VGRF, couple moment, COPl, BWm, foot length & width

5

6

25%

100

Optimized tendon force & tibia kinematics, cadaveric GRF & COP

HSSc

Tibial kinematics, GRF

9

6

NAo

25

Ankle, subtalar, and talonavicular joint kinematics

KULd

Horizontal & sagittal rotational kinematics, GRF, vertical tibia kinematics

6

3

0.8

100

Position of medial malleolus, lateral malleolus, metatarsal head, and heel

MCe

Tendon force, tibial kinematics

6

3

20

100

Foot bone motion

PSUf

Tendon force, tibial kinematics

6

3

12

100

Foot bone motion

SJTU

Tendon force, tibial kinematics, tibia loading force

4

6

5

100

Foot bone motion

TUHh

Tendon force, tibial kinematics

6

3

60

NAp

Plantar pressure, talonavicular and calcaneocuboid joint pressure

USi

Tendon force, vertical tibia force, anterior tibial translation

8

3

2

50

Foot bone motion

DMUj

Tendon force

8

3

1.5

63

Bone strain, bone strain rate, tendon force

CC

b

g

a

VA Puget Sound Health Care System, Seattle, WA, United States [23 33]. Cleveland Clinic, Cleveland, Ohio, United States [21,22]. Hospital for Special Surgery, New York City, New York, United States [34,35,42]. d Katholieke Universiteit Leuven, Leuven, Belgium [13 17]. e Mayo Clinic, Rochester, Minnesota, United States [41,43]. f Pennsylvania State University, State College, Pennsylvania, United States [2 12,44,45]. g Shanghai Jiao Tong University, Shanghai, China [18 20]. h Tubingen University Hospital, Tu¨bingen, Germany [38 40]. i University of Salford, Salford, United Kingdom and Iowa State University, Ames Iowa, United States [36,46,47]. j Des Moines University, Des Moines Iowa, United States [37]. k Vertical ground reaction force. l Center of pressure. m Body weight. n Degrees of freedom. o Not available. Duration of stance was 1/6th in vivo speed. p Not available. Simulated body weight was 35 kg. b c

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FIGURE 22.1 Dynamic gait simulator developed at Pennsylvania State University [2 12,44,45] that utilizes a translating carriage.

Muscle loading actuator Sliding carriage

Center bar

Knee axis

Electronic inclinometer

Linear drive unit

Gearbox Servo electric motor Position encoder Kistler force plate Rsscan pressure plate

Tendon clamp

Ground loading actuator FIGURE 22.2 Dynamic gait simulator developed at Katholieke Universiteit Leuven [13 17] that utilizes a translating carriage.

flatfoot model by sectioning the peritalar soft tissue structures and setting the posterior tibialis tendon force to zero for the duration of stance phase to simulate stage 2 posterior tibialis tendon dysfunction. Other groups have sourced cadaveric specimens that had the underlying disease of interest, for example, feet from individuals with diabetes [22].

22.3

Limitations of dynamic gait simulation

There are numerous limitations of DGSs associated with the difficult task of getting a cadaveric foot and ankle specimen to replicate the intricate behavior that occurs during in vivo gait. The fact that walking is common makes us often under appreciate how dynamic and complex it is. During the brief (B0.6 seconds) stance phase of gait, the ankle moves through approximately 30 degrees [53], the ground reaction force increases to more than one body weight [53], Achilles tendon force rapidly raises to over 1000 N [50], the eight other extrinsic tendons of the foot play different

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FIGURE 22.3 Dynamic gait simulator developed at Shanghai Jiao Tong University [18 20] that utilizes a translating carriage.

FIGURE 22.4 Dynamic gait simulator from the VA Puget Sound Health Care System [23 33] which leverages a hexapod robot as a primary component.

dynamic roles in stabilizing the arch [53], controlling the hindfoot and midfoot joints, and flexing and extending the metatarsophalangeal joints. All these activities occur in concert with one another and happen rapidly (,1 second) making it extremely difficult to accurately replicate this behavior in vitro with a cadaveric specimen. To simplify this problem, almost all DGSs temporally scale (slow down) the simulation of gait. Slower systems simulate stance phase over 60 seconds [38,40], while the fastest systems to date simulate stance phase in less than 1 second [13 17] (Table 22.1). Slowing down the simulation makes the task of accurately controlling the motion and forces much easier but limits the scope of scientific questions that can be answered with DGSs. For example, the speed at which a simulation is carried out can affect the model outputs since many biological tissues exhibit viscoelastic behavior. Therefore, it is best

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Force Platform

Force Platform Brackets Y

GND

Y

X X

X

Z MIC

Y

Z

Z TIB Z

Y

Foot Mounting Device

PLA

Microscribe

X Rotopod

Rotary Tendon Actuator

X Y

Z

ROB

FIGURE 22.5 Dynamic gait simulator from the Cleveland Clinic [21,22], which leverages a hexapod robot as a primary component.

FIGURE 22.6 Dynamic gait simulator from the Hospital for Special Surgery [34,35,42], which leverages a hexapod robot as a primary component.

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Tibial Rotation Motor (Angle-Control)

Ultrasound Motion-Analysis System

To Force Cylinders Tendon Force Cables

Internal-External Rotation

(only AT and TA shown)

Pressure Measuring Platform

Ground ReactionForce

Inclination Cylinder (Angle-Control)

Translation Stage (in down position)

TibialInclination Motion Ground Reaction Force Cylinder (Force-Control)

FIGURE 22.7 Dynamic gait simulator developed by the Medical School from Tubingen University Hospital [38 40], which employs custom mechanisms to generate force and motion.

practice to study outcomes that are not highly sensitive to the rate of dynamic loading if the gait simulator temporally scales the simulation. The robotic actuators within a gait simulator might need to produce ground reaction forces up to 1000 N or more if simulating full body weight. Achieving these high loads requires powerful actuators and robust high strength loading frames. Not only is it difficult to build such a robotic platform but also the frail nature of cadaveric specimens, often with poor bone quality, means specimens can fail under normal body weight loading. For these reasons, many gait simulators scale down the target ground reaction force. Hurschler et al., for example, carried out simulations with the ground reaction force reaching a peak of 350 N [38], while other systems replicate the full body weight [2 22,41,43]. Another common simplification is the number of actuated extrinsic tendons. Actuating fewer tendons or bundling them together into functional groups [21] reduces the number of expensive actuators needed and the complexity of controlling them. While many systems actuate all nine extrinsic tendons of the foot and ankle [23 35], it is also common to see six actuators used [2 17,38 41,43] (Table 22.1). While tibial motion in vivo exhibits 6-DOF, most motion occurs in the sagittal plane. Simplifying tibia kinematics by removing, for example, medial/lateral translation can greatly reduce the complexity of a gait simulator while still

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FIGURE 22.8 Dynamic gait simulator developed by the Mayo Clinic [41,43], which employs custom mechanisms to generate force and motion.

FIGURE 22.9 Dynamic gait simulator developed by the University of Salford jointly with Iowa State University [36,46 48], which employs custom mechanisms to generate force and motion.

maintaining the major characteristics of gait. Sharkey et al.’s DGS had only 3 DOF [2 12], while other systems replicate the full 6-DOF of the tibia to ground motion [18 35]. Modeling specific foot and ankle disease and or conditions within the cadaveric specimen is another challenging aspect of dynamic gait simulation. Not only do gait simulators need to replicate gait, but also they might need to replicate

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anatomical or functional characteristics of a particular disease or condition, for example cross-over toe deformity [26]. While some studies are able to acquire a cohort of cadaveric specimens with a particular disease of interest (e.g., diabetic feet [22]), for other groups, procuring numerous feet with a particular disease is impractical if not impossible. An alternative approach is to manipulate non-diseased feet to model the desired disease or post-surgical characteristics [4,28,39,43,54 58]. Creating valid models of diseases or post-surgical conditions within cadaveric feet is another limitation of dynamic gait simulation studies. For example, replicating the structure and function of a foot and ankle following surgical intervention can be challenging because cadaveric bones and tissues do not heal. Fused bones post osteotomy can be simulated with screws, plates, or additional hardware left implanted in the foot. Some diseases are more difficult to model than others and the fidelity of the modeled disease should be considered when interpreting results.

22.4

Clinical applications of dynamic gait simulation

Cadaveric gait simulators have been employed to address a wide range of clinical questions related to foot and ankle biomechanics. Numerous initial studies are related to the design, development, validation, and initial usage of the various gait simulators [2,13 21,24,25,35,38,41,48,59,60]. These studies have reported normative data, including: gross dynamic motion [2,18,20,41], applied muscle forces [2,13,17,18,21,25,35,38,48], ground reaction forces [2,14,15,17 19,21,24,25,34,35,38,41,48,60], force-time integral [38], tibial force [2,18,19,21], plantar pressure [2,17,38,41], center of pressure [17,38,41,48,59], ligament strain [35], tendon moment arms [59], and joint kinematics [16 20,24,34,35,38,41]. In general, unless otherwise noted below, these studies did not address specific functional or clinical questions, but rather demonstrated that the simulator was functioning as expected. A few research teams have employed dynamic cadaveric gait simulators to develop sensors or other hardware. Erdemir et al. used the DGS at Pennsylvania State University to explore the limitations of a fiber optic tendon force sensor related to loading rate and cable migration [45], and subsequently, skin motion artifact [3]. Another study with the DGS compared surface mounted strain gauges to strain-gauged bone staples [10]. The DGS was additionally used to validate another fiber optic tendon force sensor that was based on fiber Bragg grating [61]. The Robotic Gait Simulator (RGS) at the VA Puget Sound Health Care System has served as the test platform to develop a passive engineering mechanism for improving a flexor digitorum longus transfer to treat flatfoot [29]. The RGS also allowed for the in situ calibration of a shear wave speed-based tendon force device [32]. Owing to the ability to rigidly fix markers to bones, perhaps the most fundamental use of cadaveric gait simulators is the exploration of foot bone kinematics and joint function. Sharkey et al. have over a series of years used the DGS to address a range of issues, including: (1) how active and passive toe flexion contributes to forefoot loading [44], (2) quantifying normal hindfoot motion [4], (3) the accuracy of skin-mounted kinematic foot models [52], (4) the ability of skin-mounted foot models to diagnose various foot pathologies [12], and (5) the concept of midtarsal joint locking [11]. Lee et al. studied the effect of diabetes on midfoot joint pressures [22]. Nester et al. explored the effect of rigid body assumptions on foot kinematics [47]. Normative foot bone kinematics have been reported by some groups [20,24] while others have considered flatfoot kinematics [43]. The effect of muscle function on foot function has been studied using DGSs. Ward et al. explored how tibial and Achilles tendon forces affected forefoot and hindfoot loading [62]. Anterior tibial tendon dysfunction was shown to alter tarsal bone motion and plantar pressure [63]. The function of the flexor hallucis longus was investigated in two studies using the DGS [6,7]; a third study from this team examined isometric extrinsic toe flexor function [5]. The moment arms of the nine extrinsic muscles that cross the ankle joint have been quantified [64]. Finally, Burg et al. demonstrated that the action of specific muscles alter foot joint kinematics [13]. Various studies have examined aspects of ligament and foot function, including spring ligament strain in intact feet and feet with posterior tibial tendon dysfunction [35], the loading of the plantar fascia during stance phase while undergoing a partial release [65], and the pressure distribution in the ankle joint [40]. The DGS was used along with a fiber optic sensor to quantify plantar aponeurosis loads during normal gait [66]. The DGS was also employed to quantify the effect of medial and lateral ankle injury and repair on subtalar joint motion [9]. Another research group from the Mayo Clinic employed a cadaveric gait simulator to explore the contribution of ankle ligaments on joint stability [67]. Gait simulators have also been used to study foot type, with flatfoot simulations being conducted by several groups [35,43,68]. One of these groups (VA Puget Sound) also used the RGS to correlate long second metatarsals with increased plantar pressure [26]. Perhaps, the most common clinical area of research that has been studied using cadaveric gait simulators is that of total ankle replacements. Research teams have used DGSs to conduct studies on foot and ankle biomechanics post-total ankle replacement. These studies found increased joint laxity [69], altered ground reaction forces [15], and modified bone kinematics [16] following total ankle arthroplasty. Another area of research related to total ankle replacements is

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the effect of a joint replacement [70] and potential misalignment of the implant components [30,33] on distal foot bone kinematics. Cadaveric gait simulators have also been used to explore a wide range of orthopedic foot and ankle pathologies and treatments. Neil Sharkey’s group quantified bone strain and microcracks of the metatarsals during simulated gait [10,71,72]. Tendon transfers, including split transfers of tibialis anterior and posterior [73], and transfer of flexor digitorum longus to treat flat foot deformity [27], have been studied using the DGS and the RGS, respectively. The effect of ankle arthrodeses on adjacent foot joint kinematics have been studied at the Hospital for Special Surgery [42]. Other studies have explored how ankle fusions can alter foot joint kinetics [39] and the effect of anterior/posterior malaligned arthrodesis on foot kinematics and plantar pressure [31]. The DGS was used to study the kinematic behavior of the ankle following malleolar fracture repair [8]. The RGS has been employed to quantify how second metatarsal shortening alters plantar pressure [28] and the optimal angle of fusion for the great toe joint [23]. Finally, the effects of foot orthoses on both posterior tibial tendon friction [74] and second metatarsal bone strain has been quantified [37].

22.5

Areas of future biomechanical research

The most obvious area of future research for cadaveric gait simulation is to address the aforementioned limitations. That is to build systems that operate at full body weight (100% GRF), physiologically realistic speeds, and with full 6DOF tibia motion. There are only a few gait simulators that currently meet those high standards. Overcoming these limitations will be achieved through improved robotic hardware, better performing software controllers, and gaining access to high quality cadaveric specimens. The second area of future work that has the potential to significantly change the utility of cadaveric gait simulators is the development of new controllers that are more biofidelic. Currently, most gait simulators are programed to replicate gait by mimicking pre-recorded tibia kinematics and GRFs and applying pre-determined muscle forces during the simulation. A much more powerful tool for scientific exploration would be gait simulators that are controlled with higher level and more abstract objective functions, such as walk at 1.0 m/s while maintaining balance. The development of controllers that produce stable, realistic walking patterns without tracking pre-recorded targets would allow the researcher to explore hypothetical “what if” scenarios without testing those scenarios on living subjects first. For example, with such a system one could ask questions such as, if we fuse these joints in the foot what is a realistic gait pattern that would evolve in response and how do the internal states in the foot (bony motion, pressures, torques, forces, etc.) respond. This line of research could also be expanded to orthoses, whereby a researcher could determine foot, ankle, and tibia kinematics and kinetics in response to different ankle-foot orthoses. One avenue to create high level biofidelic controllers is to utilize hardware in the loop computer simulations. Imagine for example, a computational musculoskeletal model with a high-level task controller such as walk at 1.0 m/s while maintaining balance. This software simulation could be coupled to the cadaveric gait simulator with each side passing kinematic and kinetic data to one another. The computational model would simulate all the dynamics and muscle actuation of the human proximal to the mid-tibia, while the cadaveric gait simulator hardware would simulate the system distal to the mid-tibia. This approach allows the physical cadaveric system to model the foot, ankle and ground interaction forces with a real cadaveric specimen, achieving a level of model complexity well beyond the current state-of-the-art in computational modeling, while at the same time the cadaveric simulation gains the advantage of simulating the musculoskeletal dynamics of the rest of the body which ultimately drives the mid-tibia interaction force. An excellent exploration into this type of approach is presented by Natsakis et al. who developed a novel inertial controller to eliminate the need to pre-record simulation set points [15]. A final area of future research is to simulate tasks that are faster and higher energy. For example, cadaveric simulation of running, jumping, or rapid turning would be scientifically very appealing. To date, these rapid movement patterns occur at speeds that are beyond the capabilities of cadaveric gait simulators. In the future, faster robotic hardware and improved control systems may allow researchers to explore these dynamic movements.

22.6

Conclusion

In conclusion, cadaveric gait simulators have a long history of helping us better understand the foot and ankle. Through hypothesis-driven experimentation, cadaveric gait simulators have enabled researchers to explore foot and ankle pathologies and gain insights into established and novel interventions. Robotically actuating the foot and ankle to recreate gait has afforded us the opportunity to learn about the foot and ankle in a manner that would not be possible through dissection alone. For hundreds of years, cadaveric specimens have contributed to scientific exploration, and even in the

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modern era, their utility remains strong. Future work will increase the usefulness of cadaveric gait simulators leading to new discoveries and ultimately better health outcomes for those ailed by lower limb disease, trauma, or dysfunction.

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[58] Fukuda T, Haddad SL, Ren Y, Zhang LQ. Impact of talar component rotation on contact pressure after total ankle arthroplasty: a cadaveric study. Foot Ankle Int 2010;31(5):404 11. [59] Kim KJ, Uchiyama E, Kitaoka HB, An KN. An in vitro study of individual ankle muscle actions on the center of pressure. Gait Posture 2003;17(2):125 31. ,https://pubmed.ncbi.nlm.nih.gov/12633772/. [accessed 17.02.22] [60] Aubin PM, Cowley MS, Ledoux WR. Gait simulation via a 6-DOF parallel robot with iterative learning control. IEEE Trans Biomed Eng 2008;55(3):1237 40. ,http://www.ncbi.nlm.nih.gov/pubmed/18334421. [accessed 23.09.20] [61] Behrmann GP, Hidler J, Mirotznik MS. Fiber optic micro sensor for the measurement of tendon forces. Biomed Eng 2012;11. ,https://pubmed. ncbi.nlm.nih.gov/23033868/. [accessed 23.02.22] [62] Ward ED, Phillips RD, Patterson PE, Werkhoven GJ. 1998 William J. Stickel Gold Award. The effects of extrinsic muscle forces on the forefoot-to-rearfoot loading relationship in vitro. Tibia and Achilles tendon. J Am Podiatr Med Assoc 1998;88(10):471 82. ,https://pubmed. ncbi.nlm.nih.gov/9791951/. [accessed 23.02.22] [63] Wu¨lker N, Hurschler C, Emmerich J. In vitro simulation of stance phase gait part II: Simulated anterior tibial tendon dysfunction and potential compensation. Foot Ankle Int 2003;24(8):623 9. ,https://pubmed.ncbi.nlm.nih.gov/12956568/. [accessed 23.02.22] [64] McCullough MBA, Ringleb SI, Arai K, Kitaoka HB, Kaufman KR. Moment arms of the ankle throughout the range of motion in three planes. Foot Ankle Int 2011;32(3):300 6. ,https://pubmed.ncbi.nlm.nih.gov/21477550/. [accessed 23.02.22] [65] Ward ED, Smith KM, Cocheba JR, Patterson PE, Phillips RD. 2003 William J. Stickel Gold Award. In vivo forces in the plantar fascia during the stance phase of gait: sequential release of the plantar fascia. J Am Podiatr Med Assoc 2003;93(6):429 42. ,https://pubmed.ncbi.nlm.nih. gov/14623987/. [accessed 23.02.22] [66] Erdemir A, Hamel AJ, Fauth AR, Piazza SJ, Sharkey NA. Dynamic loading of the plantar aponeurosis in walking. J Bone Jt Surg 2004;86 (3):546 52. ,https://pubmed.ncbi.nlm.nih.gov/14996881/. [accessed 23.02.22] [67] Watanabe K, Kitaoka HB, Berglund LJ, Zhao KD, Kaufman KR, An KN. The role of ankle ligaments and articular geometry in stabilizing the ankle. Clin Biomech 2012;27(2):189 95. ,https://pubmed.ncbi.nlm.nih.gov/22000065/. [accessed 23.02.22] [68] Jackson LTLT, Aubin PMPM, Cowley MSMS, Sangeorzan BJBJ, Ledoux WRWR A robotic cadaveric flatfoot analysis of stance phase. J Biomech Eng 2011;133(5):051005. ,http://www.ncbi.nlm.nih.gov/pubmed/21599096. [accessed 01.06.14] [69] Watanabe K, Crevoisier XM, Kitaoka HB, et al. Analysis of joint laxity after total ankle arthroplasty: cadaver study. Clin Biomech 2009;24 (8):655 60. ,https://pubmed.ncbi.nlm.nih.gov/19632017/. [accessed 23.02.22] [70] Saito GH, Sturnick DR, Ellis SJ, Deland JT, Demetracopoulos CA. Influence of tibial component position on altered kinematics following total ankle arthroplasty during simulated gait. Foot Ankle Int 2019;40(8):873 9. ,https://pubmed.ncbi.nlm.nih.gov/31244338/. [accessed 23.02.22] [71] Donahue SW, Sharkey NA. Strains in the metatarsals during the stance phase of gait: implications for stress fractures. J Bone Jt Surg 1999;81 (9):1236 44. ,https://pubmed.ncbi.nlm.nih.gov/10505520/. [accessed 23.02.22] [72] Donahue SW, Sharkey NA, Modanlou KA, Sequeira LN, Martin RB. Bone strain and microcracks at stress fracture sites in human metatarsals. Bone 2000;27(6):827 33. ,https://pubmed.ncbi.nlm.nih.gov/11113394/. [accessed 23.02.22] [73] Piazza SJ, Adamson RL, Moran MF, Sanders JO, Sharkey NA. Effects of tensioning errors in split transfers of tibialis anterior and posterior tendons. J Bone Jt Surg 2003;85(5):858 65. ,https://pubmed.ncbi.nlm.nih.gov/12728036/. [accessed 23.02.22] [74] Hirano T, McCullough MBA, Kitaoka HB, Ikoma K, Kaufman KR. Effects of foot orthoses on the work of friction of the posterior tibial tendon. Clin Biomech 2009;24(9):776 80. ,https://pubmed.ncbi.nlm.nih.gov/19700230/. [accessed 23.02.22]

Chapter 23

Finite Element Modeling Panagiotis Chatzistergos1, Sara Behforootan2, Roozbeh Naemi1 and Nachiappan Chockalingam1 1

Centre for Biomechanics and Rehabilitation Technologies, Staffordshire University, Stoke-on-Trent, United Kingdom, 2Imperial College London,

Department of Surgery and Cancer, London, United Kingdom

Abstract Finite element (FE) modeling is a technique to study the internal loading of the human body in a noninvasive manner. This unique ability of FE modeling combined with its capacity for virtual experimentation have enabled exploring aspects of foot biomechanics that cannot be investigated experimentally. The newly gained insight has helped to improve footwear design and to enhance the effectiveness of therapeutic interventions. Despite its widespread use; FE modeling is still limited within the research domain and has no direct impact on clinical practice until now. The potential clinical relevance of FE analysis can be significantly enhanced by the development of robust techniques for subject-specific modeling that can be used to inform daily clinical practice. Existing applications of FE modeling for the material characterization of plantar soft tissues, for the real time assessment of internal tissue loading, or the design optimization of footwear interventions highlight the potential use of FE modeling to enhance the management of conditions such as the diabetic foot or heel pain syndrome. However, realizing this potential will require substantial technological developments to reduce the labor intensity and enhance the reliability of FE modeling. In this context, the overarching aim of this chapter is to support clinically relevant FE modeling in the area of foot biomechanics by providing an overview of existing applications and modeling techniques and by discussing the challenges, problems, and possible solutions for reliable subject-specific, clinically applicable FE modeling. Prior knowledge of FE or computer modeling is not needed for reading this chapter, which could be equally relevant to people designing FE analyses in the area of foot biomechanics and to people who want a deeper understanding of literature in this area.

23.1

Introduction

Finite element (FE) analysis is a powerful numerical method for solving complex engineering problems. The ability of FE analysis to estimate the internal stresses of biological structures together with the capacity for performing virtual experiments and simulating loading scenarios that cannot be studied in vivo has made FE modeling one of the most popular methods in biomechanics research. However, despite its widespread use and the valuable insight it has offered, FE modeling is still limited to the research domain and has yet to make a direct impact on clinical practice. Clinically relevant subject-specific FE modeling can open the way for novel diagnostic techniques and novel methods for treatment planning, which would significantly enhance clinical practice. In this context, the overarching aim of this chapter is to support clinically relevant and clinically applicable FE modeling in the area of foot biomechanics.

23.2

Basic concepts of finite element modeling

Predicting how a solid body will deform under a known load requires solving the equations of equilibrium and compatibility for every single point (i.e., infinitesimal volume) inside this solid body. Regardless of how easy or difficult these equations are to solve for a single point, the fact that a solid body has infinite points makes this problem impossible to solve. FE analysis is a numerical (i.e., approximate) method that can transform this insolvable problem into a new one that can be solved by a computer. To achieve that, the original solid body is divided into smaller bodies (i.e., FEs) which are interconnected at points common to two or more elements; these points of connectivity are called nodes. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00007-X © 2023 Elsevier Inc. All rights reserved.

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(A)

(B)

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FIGURE 23.1 Typical tetrahedral (A, B) and hexahedral (C, D) finite elements; with corner nodes only (i.e., linear elements) (A, C) and with middle nodes too (i.e., quadratic elements) (B, D).

These elements can be 2D or 3D, and they have simple, idealized geometry (e.g., pyramids, bricks). If their mechanical properties are known, it is easy to calculate how they will deform under known loads. In this case the model is solved only at the nodes; namely for a finite number of points while the solution in any arbitrary point inside a FE is calculated by interpolation. Calculations for individual elements can then be used to assess how the entire body will be affected by the externally applied loads and to calculate its internal stresses and strains. The functions used for interpolation are called shape functions and the most common types are first order polynomials (linear) or second order polynomials (quadratic). Elements that use linear shape functions have only corner nodes while elements that use quadratic shape functions also have middle nodes (Fig. 23.1). Even though this problem can be solved, it is clear that having more elements and nodes in a model increases the number of equations that must be solved and therefore the computational power that is needed to solve them. In general, a more realistic simulation of geometry and the inclusion of fine geometrical detail increases the number of elements that are needed for meshing and the computational cost of the analysis. Other parameters that increase the computational cost of an FE model include the simulation of complex mechanical behavior (e.g., hyperelasticity, viscoelasticity) or complex boundary conditions such as contact between bodies. This creates a situation where the level of detail that can be included in an analysis has to be considered with respect to the available computational power. Characteristically, the first FE study in the area of foot biomechanics was published in the early 1980s and it included a simple 2D model of a cross-section of the foot with 342 elements [1]. Thirty five years later, similar studies presented 3D models of the entire foot with up to 400,000 elements [2,3]. Despite the exponential increase in available computational power, the complexity of biological systems is so high that assumptions and simplifications must still be made. Finding the right set of assumptions and simplifications that enable the drawing of reliable and relevant conclusions is always central to every successful FE analysis.

23.3

Applications of finite element analysis in foot biomechanics

FE studies in the area of foot biomechanics cover a wide range of applications including the investigation of foot and footwear interaction, the study of the biomechanics of the physiologic foot, the material characterization of its soft tissues, and finally the assessment of the effect of various pathological conditions.

23.3.1 Simulation of the interaction between foot and footwear Studying the interaction between the foot and footwear has been one of the most common applications of FE modeling in the area of foot biomechanics [2]. Considering the unique ability of FE modeling for estimating internal tissue

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(A)

367

(B)

Sole

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Shankpiece

Upper FIGURE 23.2 A geometrically detailed FE model of the female foot and high heel shoe (A); a plantar view of the bony structures of the model with all included ligaments (B) and muscles as identified by their insertion points (C). From Figure 1 in Yu J, Cheung JT-M, Wong DW-C, Cong Y, Zhang M. Biomechanical simulation of high-heeled shoe donning and walking. J Biomech 2013;46:2067 74.

stresses and its capacity for virtual experimentation, researchers around the world have used FE modeling to study the effect of different types of footwear or orthoses on the foot, and to improve their design and their mechanical characteristics. In this context, Yu et al. used FE modeling to assess the effect of wearing high-heel shoes on the foot [4]. For this purpose, they developed an anatomically detailed model of the foot and ankle complex coupled with a model of a high-heel shoe. The foot-ankle model was designed using magnetic resonance imaging (MRI) and comprised anatomically accurate bony structures and some major ligaments of the foot embedded into a bulk soft tissue model [5]. This model was originally used to simulate balanced standing, but it was later modified to simulate the act of wearing the shoe (i.e., shoe donning) and important instances of gait (i.e., heel strike, midstance, and push off) [6] (Fig. 23.2). These analyses offered an estimation of the internal loading of the soft tissues, bones, and joints of the foot and enabled the direct quantification of how these are affected by footwear design to potentially increase the risk for injury or for the development of deformity. These studies [4 6] highlight the potential role FE modeling can play for the optimization of footwear. One area where design optimization is very important is the provision of patient-specific therapeutic footwear and orthoses that aim to reduce or redistribute plantar loading. This type of footwear is routinely prescribed to people with rheumatoid arthritis or diabetic foot syndrome to protect their feet from loading that is painful or possibly injurious. Even though the purpose of this intervention is clear (i.e., pressure reduction), designing a shoe or an orthosis to achieve this aim is still based on the clinical experience of the attending orthotist [7]. Currently there are no established clinical guidelines as to how these interventions should be designed, and their provision is still based on the skill of the orthotist, as well as trial and error. FE modeling could significantly enhance the design process of custom footwear/orthoses and promote evidence-based prescriptions; however, its use in clinical practice is extremely limited. One of the main barriers limiting the use of FE analysis is the labor burden involved in designing complex subject-specific models. To overcome this challenge Spirka et al. developed a geometrically idealized 3D model of the forefoot which could be modified based on measurements from medical imaging to account for subject-specific morphology [7]. This simplified model of the foot comprised a combination of rigid spheres and cylinders, simulating the metatarsals embedded into a volume of compliant material simulating bulk soft tissue (Fig. 23.3A). The level of geometrical complexity that is needed to achieve satisfactory accuracy was established by comparing the results of idealized forefoot models against an anatomically detailed FE model designed from computed tomography (CT) images [8]. The feasibility of virtually optimizing the design of orthoses using this modeling technique was demonstrated in a clinical study involving people with diabetes and peripheral neuropathy [9]. All participants in this study had elevated plantar pressure in the forefoot region (peak pressures .750 kPa during barefoot walking) and required orthoses to redistribute plantar loading. Virtually optimized orthoses that had a metatarsal bar were manufactured and their pressure relieving capacity was compared against orthoses that were designed and manufactured according to standard clinical practice. For this purpose, the internal geometry of the previously described foot model [8] was calibrated with the help

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FIGURE 23.3 The geometrically idealized model of the forefoot that was used to optimize the design of insoles (A). A comparison between the peak pressures for a standard insole and for virtually optimized ones that were produced using milling or 3D printing (B). Statistically significant differences between conditions are indicated with “*.” Modified from Figures 2 and 3 in Telfer S, Erdemir A, Woodburn J, Cavanagh PR. Simplified vs geometrically accurate models of forefoot anatomy to predict plantar pressures: a finite element study. J Biomech 2016;49:289 94.

of ultrasound images of the forefoot while its external geometry was designed by scanning a physical impression of the foot in a foam box. The material properties of the bulk soft tissue model were inverse engineered based on dynamic ultrasound images during walking. These subject-specific FE models were then used to simulate shod walking in a series of virtual experiments to identify the optimum orthosis design. During this iterative process the height of the metatarsal bar was gradually increased and material under the metatarsal heads was removed in an attempt to reduce plantar pressures to within safe limits (,200 kPa [10]). The virtually optimized orthoses were manufactured using two different manufacturing techniques, namely 3D milling and 3D printing. The pressure relieving capacity of these insoles was compared against insoles that were produced according to standard clinical practice. The results of this study indicated that the virtually optimized orthoses were capable of substantially reducing forefoot pressure by redistributing plantar loading (Fig. 23.3B). Comparison between conditions showed that forefoot pressure in the virtually optimized orthoses (3D milled and 3D printed) was significantly lower compared to the standard orthoses but pressure in the midfoot was significantly higher [9]. Even though it is not clear whether the virtually optimized orthoses would actually be more effective at protecting the foot at risk from injury and the FE models tended to underestimate plantar pressures [9], this study demonstrates for the first time that FE modeling can be integrated into daily clinical practice to inform the design of patient-specific orthoses. Apart from the design, the materials that are used to manufacture therapeutic footwear and orthoses also play an important role in their ability to reduce or redistribute plantar loading. Particularly in the case of cushioning materials, careful selection of stiffness is highlighted as an imperative factor for achieving maximum pressure reduction. However, material selection is still based on empirical or anecdotal evidence, and no established guidelines currently exist to help healthcare professionals prescribing or manufacturing patient-specific orthoses/footwear to identify the most appropriate materials for their patients. To address this gap in knowledge Chatzistergos et al. developed a numerical method that can be used to investigate the cushioning properties of various insole materials on a subject-specific basis [11]. The methodology within this work encompassed the development of subject-specific FE models of the heel

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25 20 15 240N 160N 80N

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FIGURE 23.4 The effect of different heel pad stiffness (A), thickness (B) and loading (C) to the optimum insole stiffness. Insole stiffness is quantified with the help of its initial shear modulus (GInsole). The insole stiffness achieving maximum peek pressure reduction (i.e. optimum stiffness) relative to a rigid surface is indicated for each condition. Adapted from figure 7 in Chatzistergos PE, Naemi R, Chockalingam N. A method for subject-specific modeling and optimisation of the cushioning properties of insole materials used in diabetic footwear. Med Eng Phys 2015;37531 38.

pad based on ultrasound indentation. These models were used to inverse engineer the material properties of the heel pad and to simulate the contact between the plantar soft tissue and a flat insole. Once the accuracy of the FE models was validated, this modeling procedure was utilized to identify the factors that are important for material selection. Briefly, a subject-specific model of the heel was used to simulate standing on a layer of cushioning material. Starting from a very stiff material, the stiffness of the cushioning material was gradually reduced to identify the one that minimizes plantar pressure (Fig. 23.4). This process was repeated for different cases of heel pad stiffness, heel pad thickness, and for different magnitudes of loading. The results of this analysis confirmed that careful selection of cushioning material stiffness can indeed significantly improve the capacity for pressure reduction [11]. Plantar pressure was affected by changes in heel pad stiffness, thickness, and loading; however, the insole stiffness that maximized pressure reduction was not sensitive to the thickness or to the stiffness of the heel pad (Fig. 23.4). This finding indicates that people with different stiffness or thickness in their plantar soft tissue do not necessarily need different cushioning materials to minimize plantar pressure. On the contrary, optimum cushioning material stiffness depends strongly on the magnitude of loading, with stiffer materials needed in the case of higher loads to minimize pressure. Subject-specific loading and factors that might affect loading (e.g., body mass, or level and type of activity, etc.) appear to be very important for material selection [11]. The validity of these findings was tested in vivo in a cohort of healthy individuals [12]. For this purpose, a range of flat insoles was produced from ten custom cushioning foam materials, which exhibited similar mechanical behavior but with varying stiffness (from very soft to very stiff). The participants were asked to stand on and walk in these materials to assess their capacity to reduce plantar pressure. The results of this study confirmed the importance of careful cushioning material stiffness selection. Optimum cushioning material stiffness achieved, on average, 16% and 40% higher pressure reduction compared to the next softer or the next stiffer material respectively during standing. In the case of walking, optimum cushioning material stiffness improved pressure reduction by 48% and 19% compared to the next softer or the next stiffer material respectively [12]. Moreover, the optimum stiffness is strongly correlated to body mass and body mass index of the participants, with stiffer materials needed in the case of people with larger values. These are the first studies to provide quantitative data to support the importance of stiffness optimization and set the basis for clinically applicable methods for material selection [11 13].

23.3.2 Simulation of healthy foot biomechanics The biomechanics of a normal foot has been investigated in a series of purely numerical studies or in combination with in vivo/in vitro observations and tests. These numerical studies have mainly focused on investigating the internal stresses in the foot during different phases of gait [14] or during standing [15,16], and on the relationship between foot function and foot shape [17], foot loading [18] or tissue mechanical properties [19,20]. Besides exploring aspects of foot biomechanics that cannot be studied either in vivo or in vitro these numerical studies also offered new insight on the factors that could increase the risk of injury.

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FIGURE 23.5 Comparison between the internal distribution of stress and the trabecular architecture in the calcaneus. (A) The distribution of static principal stresses in a longitudinal cross-section through the entire foot in the direction of the second ray; (B) The distribution of von Mises stresses in the calcaneus during standing; (C) the actual trabecular architecture from a cadaveric sample (sagittal section); (D) the transfer of tension (’bright) or compression (-’ dark) stresses in the calcaneus according to the FE analysis. From Figure 3 in Gefen A, Seliktar R. Comparison of the trabecular architecture and the isostatic stress flow in the human calcaneus. Med Eng Phys 2004;26:119 29.

An interesting study of the internal stresses of the calcaneus was presented by Gefen and Seliktar who combined FE modeling with in vitro observations to explore the manifestation of Wolf’s law in the development of the trabecular architecture of the calcaneus [16]. For this purpose, they used an FE model to estimate the stresses inside the calcaneus during standing and to identify the directions of the principal stresses. The numerically identified directions of the internal principal stresses were then compared against the architecture of trabecular bone in the calcaneus from cadaveric specimens (Fig. 23.5). The good agreement that was observed between the directions of principal stresses and the insertion angles of real trabecular paths suggests that trabecular architecture inside the calcaneus is mainly defined by the characteristics of standing. This study is a direct example of the potential of FE analysis for the investigation of fundamental questions that cannot be answered using purely experimental methods. Another example is the study by Sun et al. [17] which focused on the relationship between foot shape and function. The aim of that study was to assess the effect of different arch heights on other structures of the foot [17]. For this purpose, the authors designed a 3D model of the entire foot which was modified to represent feet with different arch heights. These models were used to simulate standing and to assess internal foot stresses and strains. The results indicated that increased arch height is accompanied by increased stress in the plantar fascia which could be linked to higher risk for developing pathologies like plantar fasciitis [17]. In a similar study on the effect of foot loading Cheung et al. [18] showed that the tension on the Achilles tendon is coupled with the strain of the plantar fascia, indicating that scenarios that lead to increased Achilles tendon tension (e.g., intense muscle contraction or excessive passive stretching of a tight Achilles tendon) could also overstrain and injure the plantar fascia [18]. The same model was used in a separate study to investigate the effect of plantar soft tissue stiffening in plantar pressure and internal stress [20]. The results of this study indicated that stiffening of the plantar soft tissue can reduce the tissue’s ability to uniformly distribute plantar loads. This in turn leads to increased peak plantar pressures for the same plantar loads and increased risk for injury (Fig. 23.6). Building on these results, a study by Behforootan et al. investigated the importance of the nonlinear nature of the stress strain behavior of plantar soft tissue on its ability to uniformly distribute plantar loads [21]. This study revealed

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Normal

F2 MPa

MPa +2.30e-01 +1.50e-01 +1.38e-01 +1.27e-01 +1.15e-01 +1.04e-01 +9.23e-02 +8.07e-02 +6.92e-02 +5.76e-02 +4.61e-02 +3.45e-02 +2.30e-02 +1.14e-02 +0.00e+00

+2.63e-01 +1.50e-01 +1.38e-01 +1.27e-01 +1.15e-01 +1.04e-01 +9.23e-02 +8.07e-02 +6.92e-02 +5.76e-02 +4.61e-02 +3.45e-02 +2.30e-02 +1.14e-02 +0.00e+00

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+2.91e-01 +1.50e-01 +1.38e-01 +1.27e-01 +1.15e-01 +1.04e-01 +9.23e-02 +8.07e-02 +6.92e-02 +5.76e-02 +4.61e-02 +3.45e-02 +2.30e-02 +1.14e-02 +0.00e+00

Peak 0.291MPa (C)

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FIGURE 23.6 The effect of soft tissue stiffness on plantar pressure distribution. Starting from an FE model with normative stiffness the stiffness of soft tissue is increased by a factor of two (F2), three (F3), and five (F5). From Figure 4 in Cheung JT-M, Zhang M, Leung AK-L, Fan Y-B. Threedimensional finite element analysis of the foot during standing a material sensitivity study. J Biomech 2005;38:1045 54.

that even when the overall deformability (for the same plantar load) of the tissue remains the same, changes in the tissue that lead to a more nonlinear stress strain behavior can also compromise the tissue’s ability to uniformly distribute plantar loading and increases the risk for injury [21].

23.3.3 Finite element modeling for the in vivo material characterization of soft tissues Reliable measurements of material properties of the soft tissues of the foot is very important for research and clinical practice. In the area of computational biomechanics, access to accurate measurements of material properties is of paramount importance for the design of reliable FE models. At the same time, there is also evidence that the assessment of the mechanical properties of tissues of the foot in the clinic could potentially enhance the clinical management of various pathologies such as the diabetic foot [22] or heel pain syndrome [23]. According to conventional engineering techniques and established experimental standards, the calculation of the mechanical properties of a material involves the isolation of samples and testing them using a material testing system to measure the material’s stress strain behavior. The extremely invasive nature of these tests makes them unsuitable for studying living human tissues. In the literature there are noninvasive mechanical tests, such as indentation, one could

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(B) (A) Ultrasound probe

Calcaneus

(C) Ultrasound probe Heel pad

FIGURE 23.7 Subject-specific 3D modeling of the heel based on ultrasound imaging for the inverse engineering of heel pad material properties. The model of the heel pad was designed using one frontal and one sagittal ultrasound image of the heel (A). The calcaneus was outlined in both images and its 3D surface was reconstructed by dragging one outline along the other (B) to design the final model (C). Adapted from Figures 1 and 2 in Behforootan S, Chatzistergos PE, Chockalingam N, Naemi R. A clinically applicable non-invasive method to quantitatively assess the viscohyperelastic properties of human heel pad, implications for assessing the risk of mechanical trauma. J Mech Behav Biomed Mater 2017;68 287 95.

use. However, these tests cannot assess the tissue’s stress strain behavior and their results depend on the shape and the dimensions of the tested tissue as well as the specific testing set-up that was used [11,21,24]. This limitation can be overcome with the combined use of noninvasive mechanical testing and FE modeling in an inverse FE analysis. In a direct use of FE modeling, the researcher assigns material properties to the different tissues of their model and then predicts their response to an external stimulus. However, in an inverse FE analysis, the material properties of the tissue are not known but its response to a specific external stimulus (e.g., indentation, compression) is known. In this case, an FE model of the noninvasive test is designed and used to predict its outcome measure (e.g., indentation force deformation curve). An optimization algorithm is then used to calculate the values of the mechanical properties of the tissue that minimize the difference between in vivo and numerical results [11,21,24 26]. To this end, Behfoootan et al. used a custom made device for ultrasound indentation to assess in vivo the macroscopic mechanical behavior of the heel pad [21]. In this case a linear array ultrasound probe was used to load the heel pad and to measure its deformations during testing. The applied loading was assessed using a load cell in series with the ultrasound probe to calculate the force-deformation curve of the indentation test. This system was used to perform two different types of tests: (1) quasi-static indentation and (2) stress relaxation to assess the elastic and viscous aspects of the mechanical behavior of the heel pad, respectively. Subject-specific 3D FE models of the indentation tests were designed based on the recorded ultrasound images and were used to numerically predict the in vivo force-deformation graphs. An optimization algorithm was utilized to calculate the values of the material coefficients that minimize the difference between the numerical and in vivo force-deformation curves of the indentation test [21]. In this study, the complex internal structure of plantar soft tissue was not considered, and heel pad was simulated using a single bulk material with subjectspecific thickness and geometry (Fig. 23.7). This limitation was addressed by Petre et al. who used a custom made loading device to compress the forefoot inside an MRI scanner [26]. In this study, subject-specific 3D models of the forefoot comprising rigid bones and layers of skin, fat, and muscle tissue were designed from MRI images. The internal deformations of these layers were assessed from load-bearing MRI images and used to inverse engineer the hyperelastic material properties of the plantar soft tissue at the forefoot [26]. The viscoelastic nature of plantar soft tissue was not considered in this case [26].

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23.3.4 Simulation of pathological conditions In the broader area of foot pathologies, FE modeling has been used to shed light on the phenomena that lead to injury, to investigate the etiology and the implications of various pathological conditions on foot biomechanics, and to assess the effectiveness of conservative and surgical interventions.

23.3.4.1 Biomechanics of the pathologic foot The ability of FE modeling for virtual experimentation has enabled researchers working in this area to investigate how foot biomechanics are affected by pathologic changes in the foot’s structure (e.g., deformity) or by alterations in the mechanical characteristics of its tissues (e.g., plantar soft tissue stiffening). Additionally, FE modeling is used to recreate and study the conditions that lead to tissue damage. By doing so, FE modeling has provided new insight on the phenomena that lead to the development of various pathologies or increase the risk for injury, which are key prerequisites for reliable early diagnosis and effective prevention. To provide a sense of the versatility and importance of FE modeling in this area two distinct applications will be discussed. The first relates to diabetic foot ulcers and the study of their etiology and the second application relates to high energy injuries and the factors that affect the risk of bone fracture. Plantar pressure ulceration in people with diabetic foot syndrome is among the main pathological conditions that have been studied using FE modeling. Ulceration in general is a complex and multifactorial complication of diabetes. However, in the case of diabetic plantar ulcers there is agreement in literature that ulceration is often caused by mechanical damage in the soft tissues of the foot that goes unnoticed because of the loss of protective sensation in the feet (i.e., peripheral neuropathy). The mechanical nature of the phenomena that trigger the development of plantar ulcers highlights the need for a deep understanding of the internal loading of the soft tissues of the diabetic foot and of the ways internal loading is affected by diabetes. In vivo biomechanical studies in age-matched groups of people with diabetes and nondiabetic participants have indicated that the plantar soft tissue of people with diabetes tends to be stiffer [27,28] and thicker compared to their nondiabetic counterparts but their epidermal plantar skin tends to be thinner [27]. The implications of these changes for the plantar soft tissue’s capacity to uniformly distribute plantar loading have been extensively investigated using models of varying complexity. These studies have shown that plantar soft tissue stiffening and skin thinning can lead to increased plantar pressure for the same magnitude of plantar load [16,20,29 31]. Assuming that the tissue’s strength (i.e., the maximum magnitude of loading the tissue can withstand without sustaining any damage) remains the same, these findings indicate that diabetic plantar soft tissue is inherently more vulnerable to trauma, compared to nondiabetic tissue, simply because its ability to uniformly distribute plantar loads is compromised. Investigation of the internal loading of diabetic plantar soft tissue has shown that internal stresses are more sensitive to changes in tissue stiffness than plantar pressure and that stresses are significantly higher closer to bony prominences compared to the surface of the foot [29,32]. Based on these findings it can be concluded that the process of ulceration is likely to start in deeper subcutaneous layers and not on the surface of the skin [15]. For a more detailed description of numerical studies on diabetic foot biomechanics the reader is encouraged to read the relevant systematic reviews by Telfer et al. [33] and Behforootan et al. [2]. Another area where FE modeling has been used to study the mechanisms of injury is high energy trauma. In this case, FE modeling is used to recreate impact loading scenarios that are impossible to capture in vivo and to study their effect on the human body. Specific areas of application include understanding the factors that increase the risk of fracture after a fall in older individuals [34] and improving road safety [35,36]. To this end, Wong et al. developed an anatomically detailed 3D model of the foot and ankle complex which was subjected to axial compressive impact loading to simulate landing on a hard surface [34]. Two different mechanical failure criteria (von Mises and Tresca) were considered to identify the areas of the bone that are vulnerable to fracture. The results of this analysis indicated that when the foot lands flat on the impact surface, the dominant type of injury is a fracture of the calcaneus which fails mainly due to excessive shear loading [34]. A parametric investigation of the effect of loading speed was used to provide insight on thresholds for possibly injurious loading. Similarly, Smolen et al. [36] investigated the effect of foot posture on the vulnerability to fracture during impact loading (Fig. 23.8). They found that the minimum impact load that can cause a fracture significantly decreased when foot posture deviated from neutral. Based on these results they concluded that the foot is least vulnerable to fracture under impact when it is in a neutral posture and most vulnerable when it is in combined eversion and external rotation [36]. The new insight these studies provide on the mechanisms of high energy trauma can inform postural guidelines for the automotive industry to minimize the risk for injury and support the design of effective protective measures and devices.

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FIGURE 23.8 The FE model that was used to study the vulnerability of the lower leg to fracture during impact loading and the effect of ankle posture; neutral (A), 18 degrees of inversion with 10 degrees of external rotation (B), 18 degrees of eversion with 10 degrees of external rotation (C), 13 degrees of dorsiflexion with 18 degrees of inversion (C) and 22 degrees of plantarflexion (E). Adapted from Figure 3 in Smolen SC, Quenneville CE. A finite element model of the foot/ankle to evaluate injury risk in various postures. Ann Biomed Eng 2017;45:1993 2008.

23.3.4.2 Study of surgical interventions Surgical interventions aim to restore the physiologic function of the injured or pathologic foot. However, due to its complex structure and function, it is not uncommon for an intervention to correct one abnormality in one part of the foot only to cause a new one in another, previously unaffected, part of the foot or other part of the body. Indeed, abnormal changes in foot biomechanics post-operatively can contribute to iatrogenic complications including joint arthritis, foot pain, etc. The effectiveness and potential adverse effects of surgical interventions are traditionally assessed in clinical studies and in biomechanical studies where the pre- and post-surgical biomechanics of the foot are compared. FE modeling, with its ability to estimate internal loading, including the contact pressure between the numerous joints of the foot, can complement traditional clinical and experimental techniques to improve the safety and effectiveness of surgical interventions. FE studies in this area tend to be problem-specific and even though they include models of varying complexity, they simulate specific aspects of a pathology or injury and do not include characteristics of individual patients [37]. The interventions that have been investigated so far include the surgical correction of deformities [38 44] and the treatment of joint degeneration [45 47], heel pain [48 51], bone fracture, or bone instability [52 55]. A detailed description of the literature on this area can be found in the relevant systematic review by Wang et al. [37].

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Modeling strategies

The exact process of performing an FE analysis might differ depending on the specific FE software that is being used but in broad terms it can be divided into three distinctive steps: (1) preprocessing, (2) solution, and (3) postprocessing. During preprocessing, the FE model is designed by defining its geometry and assigning material properties. The model is then meshed, and appropriate supports and loads (i.e., boundary conditions) are applied. Once the FE model is complete the simulation is run with the help of an appropriate solver algorithm. Even though no human input is needed during the solution process itself, the user still has to be able to select the most appropriate solution strategy for each specific application and to configure the solver for better computational efficiency. To achieve that, a good understanding of the different types of analyses and solution strategies is needed. Finally, if the solution is successfully completed, postprocessing allows the results of the simulation to be reviewed and specific measurements and calculations to be performed. In this section, strategies for key steps toward the completion of an FE analysis, namely geometry design, meshing, assignment of material properties, solver selection, and reliability assessment will be discussed separately. Before going into details for individual modeling techniques it is important to note that it is highly unlikely that a single modeling technique will be suitable for answering a range of different questions. The most appropriate methodological approach must always be decided on an application-specific basis considering the level of accuracy that is needed to draw clinically relevant conclusions. The only way to determine this is through rigorous validation. Validation is an integral part of the process of developing reliable and clinically relevant FE models, for this reason different types of validation and techniques for performing model validation will also be discussed in a separate sub-section.

23.4.1 Geometry design Designing the geometry of an FE model starts with identifying the modeling strategy that will allow the question at hand to be answered with maximum accuracy and minimum computational cost. During this process it is important to remember that designing a model that is more complex than what is needed can be as damaging as having a model that is too simplistic. To strike the right balance between complexity and computational efficiency, the researcher has to develop a very good understanding of the biomechanical and computational nature of the problem and make a series of informed decisions about the direction of their modeling approach. First, the researcher needs to decide whether the phenomenon they want to simulate takes place in 3D space or whether it could be assumed to be limited within a single plane. The second is to understand whether the specific application involves the entire foot or only specific parts of it. Finally, the researcher needs to understand the level of geometrical detail that is required for the model.

23.4.1.1 3D modeling versus 2D modeling All problems in biomechanics are inherently 3D. However, in specific circumstances assuming that a phenomenon takes place on a single well-defined plane can be a valid assumption to make. Studies that have used 2D models have focused on specific cross-sections of the foot on which tissue geometry was reconstructed from a single 2D image. 2D models of the foot were either axisymmetric or they assumed plain strain or plain strain with thickness [2]. As one can easily understand, 2D models can be substantially less labor intensive and less computationally expensive compared to 3D models. For this reason, they are particularly useful in applications where subject-specific modeling is required in big cohorts or for clinical applications. A good example is the study by Yarnitzky et al. [56] which presents a method for subject-specific modeling of the heel for the real-time calculation of its internal stresses and strains. The 2D model in this study was designed based on measurements of the curvature of the calcaneus and heelpad thickness from a single X-ray image. Subject-specific loading was calculated using an analytic model of the entire foot and was imposed to the FE model of the heel for the calculation of internal stresses and strains. The accuracy of this modeling technique was directly assessed using a synthetic foot model, and the difference between the numerically predicated and experimentally measured internal stresses ranged between 6.3% and 17% [56]. The ability to measure reliably and in real-time internal stresses would, among other things, significantly enhance the early diagnosis and prevention of diabetic foot ulcers. However, based on available evidence, it is very difficult to evaluate whether the level of accuracy reported by Yarnitzky et al. for the synthetic foot would be satisfactory for the intended use of this model. Clearly more research is still needed for the development of reliable, clinically applicable, real-time measurements of internal stresses and strain. Nevertheless, this study offers a very good example of the role FE modeling could play within a broader computational framework to transform clinical management of foot related pathologies.

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Velcro straps

(A)

subject platform

foot plate

indenter

ultrasound transducer linear actuator

force transducer

indenter displacement

(B) rigid indenter 12.7 mm

frictionless contact heel pad

t

30.0 mm axis of symmetry FIGURE 23.9 The force-controlled system for performing indentation ultrasound tests at the heel (A) and the axisymmetric FE model that was used to inverse-engineer the mechanical properties of the heel pad (B). The thickness of the heel pad model that was defined based on the measured thickness for minimal compression (#2 N). Adapted from Figures 1 and 3 in Erdemir SA, Viveiros ML, Ulbrecht JS, Cavanagh PR. An inverse finiteelement model of heel-pad indentation. J Biomech 2006;39:1279 1.

Another application where 2D modeling has been successfully used is the material characterization of plantar soft tissue. To this end, Erdemir et al. [24] used an axisymmetric model to simulate indentation tests at the heel and to inverse-engineer the hyperelastic material coefficients of the heel pad of 20 people with diabetes and 20 nondiabetic participants. In this study the indenter was a cylindrical ultrasound probe and the models were made subject-specific by using measurements of heel-pad thickness from each participant (Fig. 23.9). To enhance the subject-specific nature of the inverse-engineering process Chatzistergos et al. replaced the cylindrical probe with a linear-array one [11]. In this case the geometry of the model was reconstructed from B-mode images to generate subject-specific models of a 2D slice of the heel pad [11]. The greatest disadvantage of 2D modeling is the substantial limitations with regards to foot function and the range of loading scenarios that can be studied. More specifically simulations have to be restricted to scenarios where it is valid to assume that there are no out-of-plane forces or displacements. Moreover, if joints are simulated it must be assumed that they only rotate around a single axis that is perpendicular to the modeling plane. The impact of these limitations on the reliability of the study can, however, be reduced by focusing on specific areas of the foot and specific loading scenarios for which out-of-plane movement/loading is minimal [57].

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23.4.1.2 Modeling of the entire foot compared to anatomically focused modeling Focusing an FE model on a specific part of the foot can substantially reduce the labor intensity of geometry reconstruction and reduce the model’s computational demand. These savings in computational resources could then be used to enhance the complexity and fidelity of other aspects of the model. At the same time, not simulating the entire foot clearly restricts the scenarios that can be studied. Models that are focused on specific parts of the foot have mainly been used to study the internal loading in the plantar soft tissues of the forefoot or hindfoot [2]. These studies have also focused on loading scenarios for which it is reasonable to assume that the internal loading of these tissue would not be significantly influenced by other regions of the foot. A good example of a geometrically focused model of the forefoot can be found in the study by Budhabhatti et al. [58]. The aim of this study was to compare the effectiveness of different interventions for pathologies of the first ray of the foot. For this purpose, a 3D model of the first ray was designed from MRI images and was used to simulate the instance of gait where plantar loading of this area is largest, namely push off [58]. According to standard practice, this model was designed using conventional MRI images which are recorded under nonweight-bearing conditions. This means that the orientation of the reconstructed bone models would be different to their orientation during push off. To address this problem Budhabhatti et al. used an optimization-based technique where the orientation of the entire model and the alignment of individual bones was adjusted to minimize the difference between numerically predicted and experimentally measured plantar pressure [58]. In this case, the use of a geometrically detailed model of the entire foot would have made this highly iterative process substantially more challenging to implement. In a similar study on the hindfoot, Fontanela et al. [59] aimed to assess the effect of footwear material on the internal loading of the heel pad. For this purpose, they developed an anatomically accurate 3D model of the heel from MRI images (Fig. 23.10). This model was used to simulate heel strike, namely the phase of gait where loading in the modeled area is most intense. Their results verified the importance of cushioning material selection in reducing the internal loading in the heel.

FIGURE 23.10 A geometrically focused model for the simulation of heel strike in barefoot and shod conditions. A longitudinal section of the heel and footwear model showing the different tissue layers and materials that were considered in the analysis (A). The final FE model for barefoot heel strike (B). From Figure 3 in Fontanella SCG, Forestiero A, Carniel EL, Natali AN. Analysis of heel pad tissues mechanics at the heel strike in bare and shod conditions. Med Eng Phys 2013;35:441 7.

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23.4.1.3 Anatomically detailed compared to idealized modeling Designing an anatomically detailed FE model involves the segmentation and geometric reconstruction of individual tissues from medical imaging. On the contrary, idealized modeling of the foot involves replacing the actual geometry of individual tissues with geometrical models that are made from combinations of basic geometrical shapes such as cylinders, spheres, etc. (Fig. 23.3). In this case, medical imaging is used to measure specific geometrical features of the modeled structures and to calibrate the orientation and size of these basic shapes [7 9]. Anatomically detailed FE modeling is necessary for the study of internal loading of the tissues of the foot and therefore most of the relevant literature includes some type of geometrically detailed model [2]. However, in applications where the macroscopic behavior of the foot is more important than its internal loading and especially in cases where there is a need for subject-specific modeling in large cohorts, idealized modeling could be a better option [7 9]. Early approaches for the design of geometrically detailed models were based on X-ray images but currently almost all geometrically detailed models are designed based on images from CT or MRI [2]. CT imaging is more suitable for reconstructing the geometry of bones and MRI is better suited for the reconstruction of soft tissues. For this reason, in studies where accurate simulation of bone and soft tissue geometry was considered to be very important, images from both imaging modalities were used [35,42,60,61]. In a very small number of cases geometrically detailed models of the heel pad were also designed from ultrasound images [11,21,62]. The reconstruction of the 3D geometry of a tissue requires the identification and segmentation of the boundaries of this tissue in a series or stack of images. These images correspond to different cross-sections of the tissue on consecutive imaging planes that are parallel to one another. In the vast majority of studies this process is performed manually with the help of specialized software. Because this process can be labor intensive, anatomically detailed models are built based on imaging data from a single individual and then used as population-specific models. These models can be modified to simulate various pathological conditions, but this modeling approach cannot be used in applications where subject-specific modeling is required. To address this problem, methods for the semi-automated segmentation of foot images have been presented in the literature. However, these methods are limited to the reconstruction of the geometry of bones from CT images only; and while aided by automation, still require substantial user input [63,64]. A different solution to the same problem involves the generation of a geometrical atlas of the foot which is then adapted to match the geometry of individual subjects [65,66]. This technique is not limited to specific tissues or imaging modalities and could potentially simplify the process of developing anatomically detailed subject-specific FE models. A proposed method to achieve this is the generation of parametric models of the foot using principal component analysis [67]. This technique still requires some level of image analysis, but it does not need the segmentation of numerous CT/MRI images and therefore can substantially enhance the use of subject-specific modeling. Even though significant progress has been achieved, clearly more research is still needed before an adequately reliable method for the automated design of geometrically detailed subject-specific FE models of the foot can be achieved. Until automated geometric model generation is achieved, the most viable approach for subject-specific modeling that can be used in big cohorts remains the use of idealized models of the foot [7 9]. An exception is the application of FE modeling for the material characterization of plantar soft tissue from indentation tests. In this case, it is possible to reconstruct the 3D shape of the loaded tissue in a way that can be fully automated [2]. More specifically, in a study by Behforootan et al. [21] one frontal and one sagittal ultrasound image of the area at the apex of the calcaneus was used to outline the boundary between soft tissue and bone. Then the 3D shape of the area that was loaded during the indentation test was recreated by dragging the first outline along the second (Fig. 23.7). The subject-specific models that were created using this technique were used to inverse engineer the visco-hyperelastic material properties of the heel pad [21]. In this study, the design of subject-specific models was performed with minimum input from the operator who had to manually transfer data between different processes and perform various visual inspections and (if needed) manual corrections. However, the high contrast in ultrasound of the boundary between bone and soft tissue and the small number of images that need to be analyzed indicates that this method can be fully automated. Clearly this method does not address the problem of simplifying the generation of anatomically detailed subject-specific models for the entire foot. However, it highlights the fact that for applications where detailed representation of the form and function of the entire foot is not necessary, focusing on a small area of the foot can open the way for geometrically detailed subject-specific modeling that can be automated and used in big cohorts. In the foot, as in any other musculoskeletal structure, form and function are closely linked. Decisions about the geometry of the model can significantly affect its ability to simulate different aspects of foot function. With regards to

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FIGURE 23.11 Comparison between two models with very different descriptions of joint function; (A) geometrically detailed model that was used to study the effectiveness of different surgical interventions for the treatment of clawed hallux deformity and (C) a simplified one used to study the effect wearing socks has on plantar soft tissue loading. The internal structure of the two models is also shown (B, D respectively). Credit: Adapted from Figure 1 in Isvilanonda SV, Dengler E, Iaquinto JM, Sangeorzan BJ, Ledoux WR. Finite element analysis of the foot: model validation and comparison between two common treatments of the clawed hallux deformity. Clin Biomech (Bristol, Avon) 2012;27:837 44; and from Figure 1 in Dai XQQ, Li Y, Zhang M, Cheung JT-MM. Effect of sock on biomechanical responses of foot during walking. Clin Biomech (Bristol, Avon) 2006;21:314 21.

joint function there are two main modeling approaches used in the literature, namely detailed and simplified. A detailed description of joint function requires the reconstruction of individual bones as separate volumes and the simulation of contact conditions between them. In a more simplified approach, however, the bones of the foot are simulated as a continuous assembly of solid volumes which are linked with a flexible material to enable some relative movement at the joint level [68]. As one could easily understand, a detailed description of joint function is substantially more demanding. In this case, accurate definition of joint geometry from medical imaging is very important, while the use of contact conditions substantially increases the computational cost of the analysis. However, detailed simulation of joint function has to be included in models that are used to study internal joint loading or for applications where joint function is important. A good example of such a study is the work by Isvilanonda et al. [42] which aimed to assess joint angle correction after different surgical interventions for the treatment of clawed hallux deformity (Fig. 23.11A and B). At the same time, the simplified simulation of joint function is reliable for studies where assessing joint loading is not needed and for which simulating the overall bending stiffness of the foot is more important that the function of individual joints. A good example is the study by Dai et al. [68] which aimed to assess the effect wearing socks has on plantar soft tissue loading (Fig. 23.11C and D).

23.4.2 Meshing reating a good quality mesh is of paramount importance for a reliable FE analysis. Two parameters are particularly important: the type of element that will be used and the density of the mesh that will be created.

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23.4.2.1 Element type selection The selection of an appropriate element type should be based on the function of the simulated tissue and should be supported by the created geometry. For example, in cases where estimating the internal stresses in tendons or ligaments is not important, but simulating the transfer of forces through these tissues is important, then simple 2-node truss elements should be the obvious option [6,8,42]. Similarly, in cases where a part of the model (e.g., bones) is considered to be rigid then, instead of meshing the entire volume, meshing only its external surfaces with properly constrained planar elements will have the same effect but with considerably lower computational cost [21]. In the majority of cases, however, calculating internal stresses/strains is needed, which means that elements that can mesh areas (in 2D models) or volumes (in 3D models) are required. More specifically, in 3D models the elements that are most commonly used are tetrahedral or hexahedral elements, which can either be linear or quadratic (Fig. 23.1). The use of tetrahedral elements enables meshing even the most complex geometries with minimum input from the researcher. Meshing with hexahedral elements is significantly less automated and, in many cases, it is impossible without substantial modifications to the model’s geometry. However, despite the labor intensive nature of hexahedral meshing, this approach is considered to offer superior overall performance and to achieve higher accuracy with lower computational cost. This was verified in a comparison between different types elements specifically for foot and footwear applications [69]. This study also concluded that quadratic tetrahedral elements could potentially achieve the same performance as hexahedral elements but at higher computational cost [69].

23.4.2.2 Mesh convergence The results of an FE analysis can be strongly affected by the density of the mesh, for that reason identifying the minimum density that can produce accurate results is an important part of any successful FE analysis. This is achieved through a process where the density of the mesh is gradually increased, the model is solved and the change in at least one key outcome measure (e.g., pressure or stress) is recorded. When the results of the analysis are plotted over the total number of nodes in the model the overall trend in the graph should point to a convergence [70]. The mesh with the lowest density beyond which an increase in element number has little or no effect on results should be adopted.

23.4.3 Material properties Before assigning material properties, the researcher needs to decide for individual tissues in the model whether they need to be simulated: (1) as nonlinearly elastic or linearly elastic, (2) purely elastic or viscoelastic, (3) isotropic or anisotropic, and finally (4) whether subject-specific or population-specific properties should be used. These questions will be discussed separately for different tissues of the foot.

23.4.3.1 Bone and cartilage In the vast majority of foot related FE studies, bone and cartilage are simulated as purely linearly elastic and isotropic using population-specific material properties from literature [2]. However, in applications where the focus is on the soft tissues of the foot and the internal stresses/strains of bones and cartilage are very small or simply irrelevant then, bone and cartilage can also be simulated as rigid shells. On the contrary, when joint contact stresses are investigated then an accurate simulation of the complex mechanical behavior of cartilage is very important [71]. Similarly, in cases where highly dynamic loading scenarios or the mechanisms of impact injury are simulated then the viscoelastic nature of both bone and cartilage should also be considered [72].

23.4.3.2 Ligaments and tendons Ligaments and tendons exhibit nonlinear, viscoelastic, and anisotropic mechanical behavior. Their complex mechanical behavior has been simulated, with different degrees of fidelity, in studies investigating ligament/tendon biomechanics or pathomechanics and the effectiveness of relevant interventions [2]. Accurate simulation of the time dependent nature of their mechanical behavior is also very important for the reliable prediction of foot function during dynamic loading such as gait or impact [72,73]. In most studies in this area tendons and ligaments were modeled as linearly elastic; however, more detailed simulations of their complex mechanical behavior can also be found [42,60,61,74]. It should be noted that the available information on ligament or tendon mechanical properties is relatively limited. Even though assessing the in vivo mechanical behavior of some tendons is feasible [75] calculating subject-specific

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material properties for all tendons and ligaments of the foot remains practically impossible. For this reason, in the majority of cases population-specific material properties from a combination of literature sources were used.

23.4.3.3 Soft tissues The soft tissues of the foot include skin, adipose tissue (i.e., fat), and muscle. However, the majority of foot-related FE studies has considered some type of bulk tissue as a combination of two or even all three tissues together. This type of simplification has been widely used in studies where the key output measure is plantar pressure. However, a detailed description of the internal structure and mechanical behavior of the soft tissues of the foot is very important when the focus is on their internal stress/strain distribution and on the investigation of the mechanisms of soft tissue injury. On the contrary, based on existing literature it can be said that detailed simulation of soft tissue structure and biomechanics is not particularly important in the case of studies focusing on bone fracture fixation and the biomechanics of ligaments and tendons. In these particular cases, the simulation of soft tissues using a single bulk linearly elastic material appears to offer satisfactory accuracy [2]. When simulation of the nonlinear mechanical behavior of adipose or bulk soft tissue is needed, the use of the Ogden hyperelastic model (first order) appears to be capable of accurately representing the tissue’s response to loading with the minimum number of material coefficients [21]. In the case of skin, a range of different material models has been used including Ogden, Neo-Hookean, and Jamus Green-Simpson hyperelasticity [2]. Skin has also been simulated as a fiber-reinforced anisotropic hyperelastic material [60,61,74]. This type of modeling is particularly relevant in studies aiming to investigate injury mechanisms in skin [60]. Finally, muscle tissue has been simulated as a separate tissue in a relatively small number of studies using the Ogden, Mooney Rivlin, or neo-Hookean hyperelasticity models [2]. Similar to before, the time-dependent nature of the mechanical behavior of the soft tissues of the foot becomes very important in the simulation of dynamic loading scenarios such as gait and impact loading [35,72,73,76]. One of the main differences of soft tissues of the foot compared to all other tissues discussed here is that their subject-specific material characterization is possible. The use of subject-specific properties appears to be very important for the real time assessment of internal tissue loading and for assessing the risk for soft tissue injury. Predictions of plantar pressure are sensitive to soft tissue material properties [77] indicating that the use of subject-specific properties would enhance the reliability of studies in this area. On the contrary, subject-specific material properties have been proven not to be important for optimizing the cushioning properties of footwear or orthoses [11].

23.4.4 Solver selection FE problems can be solved using two different approaches: implicit or explicit. Linear, time-independent problems can be solved using implicit methods while nonlinear or time-dependent problems can be solved using either approach. An analysis becomes nonlinear either because of a nonlinearity in material properties (e.g., inclusion of hyperelastic materials) or in geometry (e.g., large rotations, buckling) or due to nonlinear boundary conditions (e.g., simulation of contact conditions). On the other hand, an analysis is time-dependent when the effect of acceleration is significant and cannot be ignored. Implicit solvers are iterative, incremental, and unconditionally stable. This means that the solution is achieved through a repetitive process that tries to satisfy the conditions of equilibrium. If the problem is linear and timeindependent, then the FE model is solved in a single loading step. However, if the problem is nonlinear or timedependent, then the overall applied loads or displacements are divided into smaller steps where the solution of each step is based on the solution of the previous ones. Because implicit solvers are unconditionally stable, if they converge then they always converge to the correct solution; however, convergence is not always possible. If for some reason the solver is not able to converge, then it splits the loading step into smaller increments and tries again. One of the main limitations of implicit solvers is that the successive reduction in increment size drastically increases the computational cost of the analysis and can even cause the analysis to fail. The most common reasons that could lead to a failure to converge are errors in model implementation, that could for example lead to an unconstrained model, or the generation of high deformations that lead to excessive element distortion. Usually the incrementation of the applied loads/displacement is adjusted automatically during the solution process based on the ease/difficulty to reach convergence. However, a good estimation by the user of the size of the initial loading step could reduce the number of needed adjustments and significantly improves the computational efficiency of the analysis.

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Explicit solvers are also incremental but are not iterative and are only conditionally stable. In this case the solution is divided into increments of time with each increment being solved separately. Because the solver is only conditionally stable, the solution process continuous even if the equilibrium conditions are not satisfied which in turn means that an error in an early step of the solution step will get magnified as the solution progresses. To overcome this problem the solution process has to be divided into relatively small increments. More specifically the time increment has to be smaller than the time it takes for a soundwave to cross any FE in the model [78]. This condition can lead to very small time increments [35] and substantially increase the computational cost of the analysis. As a result, explicit solvers are considered to be superior to implicit solvers mainly for the simulation of short lasting, highly dynamic phenomena such as impact [35,36,72,76,78]. On the contrary implicit solvers are considered superior for time-independent problems or for problems where the loading rate is relatively low (#0.02 m/s) [78]. Under specific conditions both approaches can face problems due to high computational demand. One way to overcome this issue is the use of parallel computing, according to which the calculations that are needed for the solution of an FE model are divided between several computers working in parallel. The efficiency of parallel computing is strongly affected by the selection of an implicit or explicit approach. In the case of implicit solvers, a load step cannot be solved unless that previous step is solved first which means that computers working in parallel must communicate a lot with each other to complete the solution. This feature of implicit parallel computing can impose a limit to the number of computational units that can be effectively used and significantly affects its computational effectiveness. On the contrary in explicit systems the part of the solution that will be assigned to individual computer units can be decoupled from the rest of the analysis which means that they are not bound by the aforementioned limitation of implicit systems.

23.4.5 Reliability assessment FE modeling is an approximate method which is based on the simplification of an inherently complex problem. Without setting proper control points in the process of model development there is a significant risk for error and the production of misleading results. Especially in the area of computational biomechanics, the complexity of biological systems makes the risk for inaccurate or misleading results even more pronounced. The only way to mitigate this risk is by ensuring the reliability of the information that is used as input in the FE model and through rigorous validation of accuracy at different stages of model development. The development of an FE model of a system as complex as the foot requires the input of numerous parameters many of which either cannot be directly measured, and are therefore assigned based on literature, or are measured with a great margin of error (e.g., material properties). Since FE models can exhibit extreme sensitivity to their input parameters in ways that would not be intuitively obvious, a good understanding of their effect on results is important. This can be achieved through sensitivity analyses where the values of uncertain parameters are changed within an acceptable range and their effect on results is directly quantified. The assessment of the overall reliability of the model can be done following two different approaches: indirect and direct validation. Indirect validation involves comparing the predictions of an FE model against relevant experimental results from the literature. On the other hand, direct validation requires the realization of an experiment that closely matches the simulated scenario and the direct comparison between numerical and experimental results. As one can understand direct validation can offer a more robust assessment of reliability than indirect validation, but it is significantly more difficult to perform. In the area of foot biomechanics, direct validation has been mainly performed by comparing numerically estimated plantar pressure measurements against in vivo measured ones for barefoot or shod conditions. A more comprehensive validation technique was presented by Tao et al. [79] who used a combination of motion capture and plantar pressure measurements. In this case, in addition to plantar pressures, the deformation of the surface of the foot during different loaded conditions that was predicted by the model was also compared against in vivo measurements. One of the greatest challenges of validation is that often FE modeling is used to estimate quantities that cannot be measured experimentally such as internal stresses and strains. A solution to this problem is the use of synthetic foot models [56]. To this end, Yarnitzky et al. developed a foot model with an anatomically accurate internal rigid skeleton embedded into a soft silicon with known mechanical properties that closely matched those of plantar soft tissue. Internal stresses in the synthetic soft tissue were directly calculated with the help of sensors that were placed under the calcaneus and between the foot model and the simulated ground [56]. Clearly, robust direct validation is very challenging, but without it the reliability and possible clinical relevance of any FE study can be severely compromised. The risk of producing erroneous and misleading results in the area of computational biomechanics cannot be overstated.

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23.5 Limitations and future research toward clinically applicable finite element modeling Despite progress in the area of subject-specific FE modeling, there are clearly some fundamental challenges that still need to be addressed before direct applications into everyday clinical practice could be considered. Among the main limiting factors are: (1) the difficulty of getting reliable input information, (2) the complexity and labor intensity of the modeling process, and (3) the inherent risk for inaccurate and misleading results. The design of an FE model requires a substantial amount of input information, mainly relating to geometry and material properties. With regards to geometry, using images from CT and MRI enables accurate reconstruction of the subject-specific geometry of the foot and should be the obvious source of geometrical information in applications where these imaging modalities are part of standard clinical practice. However, requesting either of them for the sole purpose of modeling is impractical and could significantly limit the clinical applicability of FE analysis. Ultrasound imaging on the other hand is significantly more accessible, cost-effective, and easier to use; it could support subject-specific modeling in cases where CT/MRI are not readily available, provided that the focus of the analysis is on the soft tissues close to the surface of the foot [2]. Ultrasound imaging has already been used for the subject-specific modeling of the heel [21]; however, further development of modeling processes to support modeling of the entire foot is still needed. With regards to material properties, the greatest challenge is the calculation of reliable subject-specific properties in a noninvasive and time-/cost-effective manner. The potential for subject-specific material characterization has been demonstrated for the plantar soft tissue [11,21,24,26] and some of the tendons of the foot [75] however, further work is needed to expand the range of tissues that can be assessed and to improve the clinical applicability of these techniques. Imaging modalities that offer a quantitative assessment of some aspects of the mechanical behavior of soft tissues, like ultrasound shear wave elastography, could be of great help in this direction [62]. At the same time, it is also important to recognize that getting subject-specific mechanical properties for every single tissue in the foot is practically impossible. Rigorous sensitivity analyses are required to understand when subject-specific material properties of specific tissues are important. The second major challenge is the labor-intensive nature of the whole process of FE modeling. Geometry reconstruction, in particular, remains extremely challenging. Methods to automate the reconstruction of the 3D geometry of bones from CT images have been proposed [63,64] but in the vast majority of FE analyses in the literature geometry reconstruction was performed manually. The development of image processing algorithms for the automated segmentation of different tissues from medical imaging and robust methods for the automated design of FE models are still needed. Of course, the exact level of subject-specificity of geometry must also be decided according to the specific needs of individual applications. Similar to subject-specific mechanical properties, rigorous sensitivity analyses are needed to define the optimum level of geometrical detail [8]. Finally, perhaps the greatest hurdle toward truly clinically applicable FE modeling is the reliability of results. If not properly validated, FE models can give misleading results even in the hands of highly trained and experienced users. However, to have modeling methods that can be used as part of clinical practice, then people with little or even no experience on computational mechanics should also be able to use them. There is a need therefore for modeling methods that are robust to such a degree that they can be used as black boxes. Among others, these methods will also have to have very well-defined protocols to collect the necessary input information and well-defined clinically relevant outcome measures that are easily understood by clinicians. In addition to rigorous validation of reliability of the entire process for the specific population for which it is meant to be used, the development of easy to use methods to directly assess reliability on a subject-specific basis are also needed.

23.6

Summary

FE modeling can be an invaluable tool for the study of foot and footwear biomechanics. The unique ability of FE modeling to assess the internal loading of tissues and its capacity for virtual experimentation have enabled exploring aspects of foot biomechanics that cannot be investigated experimentally. Apart from investigating the biomechanics of the physiologic foot, FE modeling has been extensively used to understand phenomena that lead to injury or the etiology of pathologies or to study the interaction between foot and footwear. The gained insight can contribute to improved footwear design and to enhanced effectiveness of therapeutic interventions (footwear and surgical). The potential clinical relevance of FE modeling has been significantly enhanced by the development of modeling techniques that open the way for subject-specific modeling in large cohorts. Existing applications of FE modeling for the material characterization of plantar soft tissues [21], the real time assessment of internal tissue loading [56], or the design optimization of

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footwear interventions [9] highlight the potential use of FE modeling to enhance the clinical management of conditions such as the diabetic foot or heel pain syndrome. However, realizing this potential will require substantial technological developments to reduce the labor intensity and enhance the reliability of FE modeling.

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[27] Chao CYL, Zheng Y-P, Cheing GLY. Epidermal thickness and biomechanical properties of plantar tissues in diabetic foot. Ultrasound Med Biol 2011;37:1029 38. [28] Klaesner JW, Hastings MK, Zou D, Lewis C, Mueller MJ. Plantar tissue stiffness in patients with diabetes mellitus and peripheral neuropathy. Arch Phys Med Rehabil 2002;83:1796 801. [29] Fernandez JW, Haque MZ, Hunter PJ, Mithraratne K. Mechanics of the foot Part 1:a continuum framework for evaluating soft tissue stiffening in the pathologic foot. Int J Numer Method Biomed Eng 2012;28:1056 70. [30] Gu Y, Li J, Ren X, Lake MJ, Zeng Y. Heel skin stiffness effect on the hind foot biomechanics during heel strike. Ski Res Technol 2010;16:291 6. [31] Thomas VJ, Patil KM, Radhakrishnan S. Three-dimensional stress analysis for the mechanics of plantar ulcers in diabetic neuropathy. Med Biol Eng Comput 2004;42:230 35. [32] Gefen A. The in vivo elastic properties of the plantar fascia during the contact phase of walking. Foot Ankle Int/Am Orthop Foot Ankle Soc Swiss Foot Ankle Soc 2003;24:238 44. [33] Telfer S, Erdemir A, Woodburn J, Cavanagh PR. What has finite element analysis taught us about diabetic foot disease and its management? A systematic review. PLoS One 2014;9:e109994. [34] Wong DW, Niu W, Wang Y, Zhang M. Finite element analysis of foot and ankle impact injury: risk evaluation of calcaneus and talus fracture. PLoS One 2016;11:e0154435. [35] Shin J, Yue N, Untaroiu CD. A finite element model of the foot and ankle for automotive impact applications. Ann Biomed Eng 2012;40:2519 31. [36] Smolen C, Quenneville CE. A finite element model of the foot/ankle to evaluate injury risk in various postures. Ann Biomed Eng 2017;45:1993 2008. [37] Wang Y, Wong DW, Zhang M. Computational models of the foot and ankle for pathomechanics and clinical applications: a review. Ann Biomed Eng 2016;44:213 21. [38] Bayod J, Becerro de Bengoa Vallejo R, Losa Iglesias ME, Doblare´ M. Stress at the second metatarsal bone after correction of hammertoe and claw toe deformity: a finite element analysis using an anatomical model. J Am Podiatr Med Assoc 2013;103:260 73. [39] Bayod J, Losa-Iglesias M, Becerro de Bengoa-Vallejo R, Prados-Frutos JC, Jules KT, Doblare´ M. Advantages and drawbacks of proximal interphalangeal joint fusion vs flexor tendon transfer in the correction of hammer and claw toe deformity. A finite-element study. J Biomech Eng 2010;132:051002. [40] Garcı´a-Gonza´lez A, Bayod J, Prados-Frutos JC, Losa-Iglesias M, Jules KT, de Bengoa-Vallejo RB, et al. Finite-element simulation of flexor digitorum longus or flexor digitorum brevis tendon transfer for the treatment of claw toe deformity. J Biomech 2009;42:1697 704. [41] Iaquinto JM, Wayne JS. Effects of surgical correction for the treatment of adult acquired flatfoot deformity: a computational investigation. J Orthop Res 2011;29:1047 54. [42] Isvilanonda V, Dengler E, Iaquinto JM, Sangeorzan BJ, Ledoux WR. Finite element analysis of the foot: model validation and comparison between two common treatments of the clawed hallux deformity. Clin Biomech (Bristol, Avon) 2012;27:837 44. [43] Matzaroglou C, Bougas P, Panagiotopoulos E, Saridis A, Karanikolas M, Kouzoudis D. Ninety-degree chevron osteotomy for correction of hallux valgus deformity: clinical data and finite element analysis. Open Orthop J 2010;4:152 6. [44] Trabelsi N, Milgrom C, Yosibash Z. Patient-specific FE analyses of metatarsal bones with inhomogeneous isotropic material properties. J Mech Behav Biomed Mater 2014;29:177 89. [45] Ozen M, Sayman O, Havitcioglu H. Modeling and stress analyses of a normal foot-ankle and a prosthetic foot-ankle complex. Acta Bioeng Biomech 2013;15:19 27. [46] Terrier A, Larrea X, Guerdat J, Crevoisier X. Development and experimental validation of a finite element model of total ankle replacement. J Biomech 2014;47:742 5. [47] Wang Y, Li Z, Wong DW, Cheng CK, Zhang M. Finite element analysis of biomechanical effects of total ankle arthroplasty on the foot. J Orthop Transl 2018;12:55 65. [48] Cheung JT, An KN, Zhang M. Consequences of partial and total plantar fascia release: a finite element study. Foot Ankle Int 2006;27: 125 32. [49] Gefen A. Stress analysis of the standing foot following surgical plantar fascia release. J Biomech 2002;35:629 37. [50] Liang J, Yang Y, Yu G, Niu W, Wang Y. Deformation and stress distribution of the human foot after plantar ligaments release: a cadaveric study and finite element analysis. Sci China Life Sci 2011;54:267 71. [51] Tao K, Ji W-T, Wang D-M, Wang C-T, Wang X. Relative contributions of plantar fascia and ligaments on the arch static stability: a finite element study. Biomed Tech 2010;55:265 71. [52] He K, Fu S, Liu S, Wang Z, Jin D. Comparisons in finite element analysis of minimally invasive, locking, and non-locking plates systems used in treating calcaneal fractures of Sanders type II and type III. Chin Med J (Engl) 2014;127:3894 901. [53] Pang Q-J, Yu X, Guo Z-H. The sustentaculum tali screw fixation for the treatment of Sanders type II calcaneal fracture: a finite element analysis. Pakistan J Med Sci 2014;30:1099 103. [54] Ramlee MH, Kadir MRA, Murali MR, Kamarul T. Biomechanical evaluation of two commonly used external fixators in the treatment of open subtalar dislocation a finite element analysis. Med Eng Phys 2014;36:1358 66. [55] Wang Y, Li Z, Zhang M. Biomechanical study of tarsometatarsal joint fusion using finite element analysis. Med Eng Phys 2014;36: 1394 400.

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[56] Yarnitzky G, Yizhar Z, Gefen A. Real-time subject-specific monitoring of internal deformations and stresses in the soft tissues of the foot: a new approach in gait analysis. J Biomech 2006;39:2673 89. [57] Erdemir A, Saucerman JJ, Lemmon D, Loppnow B, Turso B, Ulbrecht JS, et al. Local plantar pressure relief in therapeutic footwear: design guidelines from finite element models. J Biomech 2005;38:1798 806. [58] Budhabhatti SP, Erdemir A, Petre M, Sferra J, Donley B, Cavanagh PR. Finite element modeling of the first ray of the foot: a tool for the design of interventions. J Biomech Eng 2007;129:750 6. [59] Fontanella CG, Forestiero A, Carniel EL, Natali AN. Analysis of heel pad tissues mechanics at the heel strike in bare and shod conditions. Med Eng Phys 2013;35:441 7. [60] Fontanella CG, Carniel EL, Forestiero A, Natali AN. Investigation of the mechanical behaviour of the foot skin. Ski Res Technol 2014;1 8. [61] Forestiero A, Raumer A, Carniel EL, Natali AN. Investigation of the interaction phenomena between foot and insole by means of a numerical approach. Proc Inst Mech Eng Part P J Sport Eng Technol 2015;229:3 9. [62] Chatzistergos P, Behforootan S, Allan D, Naemi R, Chockalingam N. Shear wave elastography can assess the in-vivo nonlinear mechanical behavior of heel-pad. J Biomech 2018;28:114 50. [63] Camacho DLA, Ledoux WR, Rohr ES, Sangeorzan BB, Ching RP. A three-dimensional, anatomically detailed foot model: a foundation for a finite element simulation and means of quantifying foot-bone position. J Rehabil Res Dev 2002;39:401 10. [64] Chen WM, Lee SJ, Lee PVS. Strategies toward rapid generation of forefoot model incorporating realistic geometry of metatarsals encapsulated into lumped soft tissues for personalized finite element analysis. Comput Methods Biomech Biomed Engin 2017;20:1421 30. [65] Bucki M, Luboz V, Perrier A, Champion E, Diot B, Vuillerme N, et al. Clinical workflow for personalized foot pressure ulcer prevention. Med Eng Phys 2015;38:845 53. [66] Lochner SJ, Huissoon JP, Bedi SS. Development of a patient-specific anatomical foot model from structured light scan data. Comput Methods Biomech Biomed Engin 2014;17:1198 205. [67] Scarton A, Sawacha Z, Cobelli C, Li X. Toward the generation of a parametric foot model using principal component analysis: a pilot study. Med Eng Phys 2016;38:547 59. [68] Dai X-Q, Li Y, Zhang M, Cheung JT. Effect of sock on biomechanical responses of foot during walking. Clin Biomech (Bristol, Avon) 2006;21:314 21. [69] Tadepalli SC, Erdemir A, Cavanagh PR. Comparison of hexahedral and tetrahedral elements in finite element analysis of the foot and footwear. J Biomech 2011;44:2337 43. [70] Maas SA, Ellis BJ, Rawlins DS, Weiss JA. Finite element simulation of articular contact mechanics with quadratic tetrahedral elements. J Biomech 2016;49:659 67. [71] Wu L. Nonlinear finite element analysis for musculoskeletal biomechanics of medial and lateral plantar longitudinal arch of virtual Chinese human after plantar ligamentous structure failures. Clin Biomech (Bristol, Avon) 2007;22:221 9. [72] Bandak FA, Tannous RE, Toridis T. On the development of an osseo-ligamentous finite element model of the human ankle joint. Int J Solids Struct 2001;38:1681 97. [73] Wang Z, Imai K, Kido M, Ikoma K, Hirai S. A finite element model of flatfoot (pes planus) for improving surgical plan. Conf Proc IEEE Eng Med Biol Soc 2014;844 7. [74] Fontanella CG, Favaretto E, Carniel EL, Natali AN. Constitutive formulation and numerical analysis of the biomechanical behaviour of forefoot plantar soft tissue. J Eng Med 2015;228:942 51. [75] Maganaris C, Chatzistergos P, Reeves N, Narici M. Quantification of internal stress-strain fields in human tendon: unravelling the mechanisms that underlie regional tendon adaptations and mal-adaptations to mechanical loading and the effectiveness of therapeutic eccentric exercise. Front Physiol 2017;8:91. [76] Chen W, Lee PV. Explicit finite element modelling of heel pad mechanics in running: inclusion of body dynamics and application of physiological impact loads and application of physiological impact loads. Comput Methods Biomech Biomed Engin 2015;5842. [77] Akrami M, Qian Z, Zou Z, Howard D, Nester CJ, Ren L. Subject-specific finite element modelling of the human foot complex during walking: sensitivity analysis of material properties, boundary and loading conditions. Biomech Model Mechanobiol 2018;17:559 76. [78] Sun JS, Lee KH, Lee HP. Comparison of implicit and explicit finite element methods for dynamic problems. J Mater Process Technol 2000;105:110 18. [79] Tao K, Wang D, Wang C, Wang X, Liu A, Nester CJ, et al. An in vivo experimental validation of a computational model of human foot. J Bionic Eng 2009;6:387 97.

Chapter 24

Musculoskeletal Modeling of the Foot and Ankle Scott Telfer1,2,3 1

Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States, 2Department of Mechanical Engineering,

University of Washington, Seattle, WA, United States, 3RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States

Abstract The complexity of the foot and ankle’s anatomy and function makes direct measurement of its in vivo mechanics challenging. This has driven researchers to explore the use of computational simulations based on (neuro)musculoskeletal models to study foot and ankle biomechanics, an approach that has been enabled by recent developments in computational numerical techniques and accessible modeling tools. These modeling approaches have numerous potential applications, including the in silico personalization of clinical interventions, optimization of function to improve sports performance, assessment of workplace ergonomics, and more. In this chapter, the development of musculoskeletal models is described, along with the challenges involved that are pertinent to the foot and ankle. Current musculoskeletal foot and ankle models that have been reported in the literature are discussed, along with future areas of research.

24.1

Introduction

Several questions related to the mechanics and control of human motion in both healthy and disease states still need to be answered. Furthermore, the role of the foot and ankle in the overall system has received relatively little attention. As the most distal major segment and the primary interface between the body and the ground during many activities of daily living, the foot plays an essential function when performing common tasks, such as walking or running. However, the complexity of the foot’s anatomy, with 28 individual bones and over one hundred extrinsic and intrinsic muscles and ligaments, makes performing detailed studies of its function extremely difficult. Understanding the functional mechanics of any individual joint, muscle, or ligament in the body through in vivo experiments can be challenging. This has led researchers to develop and explore computational simulations using musculoskeletal models to gain insights into the biomechanics of motion. In recent years, a combination of improvements in computational techniques and the continued development of accessible tools have opened up the field of musculoskeletal modeling research. However, similar to the approaches taken in conventional gait analysis, most full-body or lower limb musculoskeletal models represent the foot as a single rigid segment, or at least require this for inverse kinetic-based analyses [1]. This significantly limits the applicability of these models in the context of studies investigating foot pathologies. While some novel, non-invasive approaches can provide direct measurements of the forces in certain tendons [2], the only real option to predict the activation patterns and forces of a large number of muscles in a complex system such as the foot and ankle is to use in silico modeling approaches. In this chapter, the development of these musculoskeletal models will be discussed, along with the challenges inherent to modeling the foot and ankle, the findings from existing studies, and future areas of research. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00021-4 © 2023 Elsevier Inc. All rights reserved.

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24.1.1 Musculoskeletal models and their development To produce movement, the central nervous system sends commands that cause the muscles of the body to contract, generating forces around joints. Various sensing cells in the body provide proprioceptive feedback about the position of these joints, giving the control system the information needed to modify its commands. In the human body, this neuromuscular system is extremely complex, and must account for many factors including the redundancy that is inherent in the musculoskeletal system (i.e., there are often multiple muscles actuating the same motion in a joint), along with contributions from ligaments and other soft tissues, as well as external forces acting on the body. An overview of how musculoskeletal models can be developed is given below. Model geometry: In general, musculoskeletal models are based upon what are assumed to be rigid body segments that represent the major parts of the body, for example, the thigh, trunk, or forearm. The overall mass of each body segment, along with the location of their center of mass and their inertial tensor needs to be known or estimated. Many current modeling platforms use bone surface models that have been segmented from computed tomography (CT) or magnetic resonance imaging (MRI) scans to allow visualization of these segments during model construction and simulations. In the model, the body segments are connected via constrained joints, often mechanical analogs, for example, a ball and socket joint for the hip, or a hinge joint for the knee. Often the number of degrees of freedom of these joints is reduced compared to the mechanics of the “real” joint for computational simplicity. For example, knee kinematics are more complex than the simple one degree of freedom hinge joint that is often used to represent it, particularly for more demanding movements where significant motion can occur outside of the primary (sagittal) plane [3]. Musculoskeletal models also require geometric information about the insertion points and paths that musculotendinous units take between and over the underlying bones. This is required so that their moment arms and lengths are accurately represented for dynamic simulations [4]. Mapping of muscle pathways can often be done via MRI, or through digitization of cadaver specimens [5]. Simulation results can be quite sensitive to the position of muscle insertion points, particularly those that play an important role in the task being studied, for example the insertion of the Achilles tendon into the calcaneus for studies on walking [6]. In reality, muscle pathways tend to be non-linear, as the muscle has to fit around the rest of the body’s anatomy, but in some cases, for modeling simplicity they are treated as straight lines. Joints with larger ranges of motion will often need to account for changes in the muscle pathways as tendons move over bony prominences when the joint is flexed or extended. For example, the flexor tendons of the first metatarsophalangeal joint “wrap” around the head of the metatarsal as the joint goes into extension. Representing this in the model is often achieved using via points or wrapping objects with simple geometry (often cylinders) that approximate the tendon passing over the bone. In models that also include ligaments, these require the same geometric insertion and pathway information as musculotendinous units. Model scaling: Given the amount of information required to recreate the geometry of a particular human in a musculoskeletal model, a process that likely requires a full body or at least limb of interest MRI scan for accurate mapping of musculotendinous geometry, the production of a complete, subject-specific model from scratch for each individual that a researcher is interested in studying is generally not feasible. To allow multiple individuals to be studied, a template model that has been previously developed and validated is often scaled to match the subject of interest. This scaling applies to all aspects of the model: mass distributions, body segment geometry, muscle pathways, etc., and can involve scaling to a combination of linear or other measurements taken from the subject being studied. Some questions do exist as to which scaling technique to use and how accurate these are when applied across a wide range of individuals. For example, the choice of scaling technique affects muscle force predictions for different muscles acting around the ankle [7], and in this case the authors recommended scaling with precalculated joint angles to improve consistency between inverse kinematic and direct kinematic approaches. Intrinsic forces (muscle modeling): The generation of force by muscles is complex and non-linear. For most musculoskeletal models, some variation on a Hill-type [8] muscle model is used, with the complete musculotendinous unit usually being represented by contractile and spring elements (Fig. 24.1). Several parameters are used to characterize each musculotendinous unit, including the maximum isometric force (often estimated relative to the cross-sectional area of the muscle), optimal fiber length, and tendon slack length, among others [9]. Several factors relating to muscle physiology are not included in Hill-type models, such as the distribution of fast and slow twitch fibers. More complex models do exist but the use of these must be balanced against computational efficiency. In addition, if ligaments are included in the model then the material properties for these also need to be available or estimated [10]. Analyses: The equations that govern the motion of the musculoskeletal system are complex, and quickly move beyond a level that can be solved manually (Fig. 24.2). It is only in recent years that advances in computational power and processing techniques have made it feasible to run musculoskeletal simulations in reasonable timeframes using

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FIGURE 24.1 Overview of a Hill-type muscle model. CE, contractile element, responsible for active force generation; SEE, serial elastic element, primarily representing the tendon; PEE: parallel elastic element, representing the connective tissue that surrounds the muscle; F: tendon force. The angle α represents the angle between the fiber direction and the tendon.

FIGURE 24.2 An example of a neural controller used for a musculoskeletal modeling simualtion, in this case using feed forward and feedback control. Credit: Kaminishia K, Chibab R, Takakusakib K, Ota J. Investigation of the effect of tonus on the change in postural control strategy using musculoskeletal simulation. Gait Posture 2020;76:298 304.

consumer-level computing hardware. Once a model has been established, there are numerous analyses that can be performed. It is beyond the scope of this chapter to cover all of these, but a few common examples are discussed below. Simulations using musculoskeletal models can generally be grouped into two types: inverse and forward dynamic simulations. Inverse dynamic simulations predict the activation patterns and forces of muscles and in some cases ligaments, usually based on experimentally captured motions and external reaction forces. These are more efficient to calculate. Forward dynamic simulations compute the kinematics of a model that are generated by the application of forces on the body (usually muscle activation patterns, gravity, and the external environment). This tends to be the more computationally expensive of the two types of simulation. Inverse dynamics: In many tasks we can often measure the kinematics of the body and external reaction forces experimentally with relatively high levels of accuracy, along with a select number of muscle activation patterns using electromyography (EMG). While standard analyses of this type of motion capture data can allow net internal or external moments about the joints to be calculated, this provides only a gross overview of the kinetics of the joint and reduces the combined interaction of all forces including multiple muscles and ligaments to a single component [11]. Gaining a

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more complete estimate of the complex muscle activation patterns generated by the central nervous system can only be achieved through the use of musculoskeletal simulations. Despite some limitations with the estimates that are produced, these modeling approaches have allowed us to better understand the roles of certain muscles and ligaments [12], joint loading patterns [13], and more. In the case of disease or deformity these results have the potential to help improve treatments, whether surgical, rehabilitative, or orthotic in nature. Techniques including static optimization [1] and computed muscle control [14] are used to estimate the contributions of each muscle to the body system at each moment in time of the task, going beyond the net moments produced in basic inverse dynamics. An inverse kinematic analysis may also be performed as part of a larger inverse dynamic analysis, and this is used to translate motion capture data, often from a set of retroreflective markers, to the best matching model pose across the timesteps of the data collection. When performing inverse dynamic simulations, there are often inconsistencies between the model and the force plate data. Tools, such as the residual reduction algorithm (RRA) [15] in OpenSim or The “Hand of God” in the AnyBody platform, are used to account for these small discrepancies. These tools attempt to track experimental data while minimizing the use of actuators that move the model. The RRA algorithm also provides recommendations to alter the mass distribution of the body to improve dynamic consistency. Forward dynamics: Forward dynamics, or predictive simulations, do not require experimentally collected kinematic or kinetic data as an input, and are often based on a high-level goal, such as the assumption that the central nervous system minimizes the metabolic cost of transport, and a task requirement like maintaining a target walking speed [16]. This approach can reveal underlying principles of movement and show cause-effect relationships. These types of simulation have the potential to be useful in optimizing clinical interventions and improving performance in sporting events, for example [17]. Early forward dynamic models were computationally expensive [18], but the adoption of optimal control theory methods such as direct collocation have allowed trajectory optimization to be performed in a more time efficient manner [19]. Using musculoskeletal models to plan treatments and predict how a proposed intervention will change an individual’s gait pattern through forward dynamic simulations is an area that is of increasing interest to researchers. A key aspect of producing realistic simulations is to accurately model foot-ground contact mechanics, the nature of which can have a large effect on predicted muscle activation patterns [20]. Studies using foot-ground contact models to reproduce ground reaction forces during walking and running have primarily used viscoelastic point-like elements placed on the bottom of the foot. Simple objects like spheres and cylinders have been used to model the shoe and plantar tissues. Researchers have placed these contact model elements in different configurations and sites across the plantar surface of the foot [21]. The parameters of the viscoelastic system are often estimated by minimizing the differences between the predicted kinematics and those measured using a training set of data [22]. Modeling platforms: While several individual research groups have produced their own custom modeling software [18], several musculoskeletal modeling platforms are available, both commercial and open source. For example, the use of OpenSim [1], and AnyBody [23] platforms are commonly reported in the literature.

24.1.2 Validation techniques The process of ensuring a musculoskeletal model produces valid results is complex and multi-factorial [24]. All models require some assumptions and approximations to be made. Models need to be carefully assessed to determine if these factors will adversely affect the simulation results. For example, many musculoskeletal models have been shown to be sensitive to certain input parameters [25]. A framework for validation and verification of neuromuscular models has been presented by Hicks et al. [24]. There are several different data sources that can be used to compare model simulation results to get a sense of model validity. EMG measurements, if not used as drivers for the simulation, can provide some insights into the timing and levels of activity of muscles that can be compared to predicted patterns (Fig. 24.3). However, the intrinsic muscles of the foot and non-surface extrinsic muscles like tibialis posterior that play an important role in foot function require invasive intramuscular EMG measurements, where an electrode is placed via needle into the muscle belly. Intramuscular EMG measurements can also provide inherently more accurate data than surface EMG but are considerably more difficult to collect. If using a musculoskeletal model to provide inputs to a finite element (FE) model, the predicted plantar pressures can be compared to experimentally measured values. Studies involving inverse kinematics based on skin mounted markers can be compared to bone pin and biplane data in the literature.

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FIGURE 24.3 Predicted muscle activation patterns (red) versus those experimentally measured via EMG for Glasgow-Maastricht foot model (blue). Note generally good agreement for muscles acting primarily in the sagittal plane and poorer agreement for those in other planes.

Revolute joints Universal joints Spherical joints

FIGURE 24.4 Representation of foot joints with restricted degrees of freedom in the Glasgow-Maastricht foot model, developed in the AnyBody modeling platform.

24.1.3 Challenges in modeling the foot and ankle Number of body segments and degrees of freedom: Given the number of bones (and joints) that make up the foot and ankle, it could potentially be divided into a large number of rigid body segments for a fully representative musculoskeletal model. Therefore, there are potentially many degrees of freedom that could be included in the model. In reality, the relative movement of most of these joints is small, for example between most of the midfoot bones, and these can often be reasonably treated as a single rigid body. In general, musculoskeletal models tend to use joint definitions with a limited number of degrees of freedom to reduce the problem space, for example, modeling the elbow as a one degree of freedom hinge. Approaches for modeling the foot are no different, for example modeling the metatarsophalangeal joints with one or two degrees of freedom or implementing kinematic rhythms (functional units) as discussed later. An example of one approach taken to restrict the number of degrees of freedom in a foot model is shown in Fig. 24.4. Anatomical data: Available data regarding the properties of the intrinsic muscles, tendons, and ligaments of the foot tends to be quite limited [26] or based on a small number of specimens [27 29]. Thus, some assumptions may be required when estimating the parameters needed to represent these in musculoskeletal models of the foot. Input kinematics: Performing inverse dynamic analyses of intrinsic foot biomechanics requires the motion of the relevant foot segments to be tracked in space. Numerous multisegment foot models have been reported in the literature with different marker sets and segments tracked [30], and researchers need to balance their choice of model with their needs and the technical limitations of their motion capture system. Because of the relatively small motion of many of the joints in the foot, tracking these motions accurately is challenging, especially when considering skin motion

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artifacts. In addition, the size of each segment, in comparison to the thigh or pelvis, often makes it difficult to physically place enough tracking markers to get accurate data from optical motion capture systems, and in the case of the talus it is not possible to use surface mounted markers. Moreover, it is not possible to accurately study the effect of footwear on bone kinematics using traditional motion analysis techniques. Biplane fluoroscopy-based bone tracking systems may overcome some of these issues; however, this is still a developing technology in the field of foot and ankle research [31]. One approach to account for the lack of experimental information about the kinematics of some foot joints is to use kinematic “rhythms,” or functional units to prescribe the motion of a group of joints [32]. For example, in the Glasgow Maastricht model the longitudinal medial arch, the plantarflexion angle of the talonavicular joint, medial cuneonavicular joint, and first tarsometatarsal joint are linked by two coefficients to form a single degree of freedom. The curvature is therefore prescribed by the talonavicular flexion angle. Input kinetics: Another important challenge relates to kinetic analyses of the joints of the foot. The majority of modern motion analysis laboratories that study walking, running, or similar movements are set up to capture ground reaction forces using force plates embedded in the floor of the lab. Each plate can provide a single force vector for each limb, which works fine if the foot is modeled as a single segment. However, when the intrinsic joints of the foot are of interest and it is divided into multiple segments, this technique is no longer valid as several segments are in direct contact with a single plate (and thus a single force vector). One relatively common approach to address this is to use a proportionality scheme, where plantar pressure measurements are collected in addition to force plate data. This can be synchronously collected, using a pressure plate mounted directly on top of a force plate to simultaneously capture regional plantar loading at the same time as the ground reaction force data [33], or asynchronously collected and combined with the force plate data during post processing. With either approach, for each instance that the foot is in contact with the plates, the ground reaction forces (vertical and shear) and moments are divided and proportionally assigned to the corresponding segments of the foot [34]. This assumes that the shear forces act uniformly across the plantar foot [35] and that peak plantar pressures and peak shear stress occur in the same location, which has also been shown to not always be the case [36].

24.2

Foot specific models and applications

A small number of musculoskeletal models which include a representation of the foot that goes beyond the typical single rigid segment have been reported in the literature. A brief summary of this research is given below. Several foot and ankle models have been developed using the AnyBody modeling platform. Initially, a threesegment foot consisting of a combined hindfoot/midfoot segment, a forefoot segment, and a toe segment was build and published in 2010 [37]. This model included 16 of the intrinsic muscles of the foot that spanned more than one of the modeled segments. It also included four ligaments that crossed between the segments of the model. The predicted muscle activation patterns from the model were compared to EMG signals reported in the literature, and in general good agreement was found between the two. A more advanced model built on the same platform, the Glasgow Maastricht foot model, was reported in 2016 and consists of 26 segments (all of the major foot bones) [32,38]. As mentioned earlier, this model includes kinematic rhythms, as well as lateral contact between the metatarsal heads and tarsal bones to reduce the overall number of degrees of freedom. Kinematically, the results were found to be comparable to those reported in bone pin studies. Kinetically, the model is designed to be driven by plantar pressure data, proportionally divided and applied to each of the segments. The Glasgow Maastricht foot model was applied in a group of five subjects to gain insights into the reaction forces of several foot joints during gait. The combined forces resulting from the mass and movement of the body combined with the actions of muscles means that the joints of the foot can have loads equivalent to several times bodyweight acting through them. This was reflected in the results from this study, which found high loads on the tarsometatarsal (peak 3.0 BW) and first metatarsophalangeal joint (peak 1.9 BW), occurring at around 80% of stance phase [39] (Fig. 24.5). A modified version of the Glasgow Maastricht foot model with five segments (Fig. 24.6) was used by Kim et al. [33] in a kinetic analyses to determine joint reaction forces for the ankle, midtarsal, tarsometatarsal, and metatarsophalangeal joints in a group of five subjects [33]. The average peak joint reaction force for the ankle was estimated to be 8.72BW, 4.31BW for the midtarsal joint, 2.65BW for the tarsometatarsal joint, and 3.41BW for the metatarsophalangeal joint. It should be noted that this study used a simple muscle model without viscoelastic components. Malaquias et al. reported on the development of foot musculoskeletal model built using the OpenSim platform that had two versions: one with 15 degrees of freedom and one with 8 degrees of freedom. This model included joints at the

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FIGURE 24.5 First metatarsophalangeal joint reaction force during stance phase for four subjects with normal foot type.

FIGURE 24.6 Musculoskeletal foot model produced using the Anybody software platform. Top: The five segments making up the model are shown along with their degrees of freedom; bottom: the intrinsic/extrinsic muscles (pink) and ligaments (white). Credit: Kim Y, Lee KM, Koo S. Joint moments and contact forces in the foot during walking. J Biomech. 2018;74:79 85.

midtarsal and tarsometatarsal joints [40]. The models also included 36 ligaments that spanned multiple segments. The inverse kinematic results from these models for five healthy subjects were compared to bone pin-derived data. The authors found that the 8-degree of freedom model had less variability and was somewhat more in line with the bone pin data, so used it for kinetic analysis. For this, they found the standard deviation of the joint torques to be small, suggesting high inter-subject consistency.

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FIGURE 24.7 The standard OpenSim foot model with all extrinsic foot muscles is shown on the left; Model 2 (center) is a recently proposed 8-degrees of freedom OpenSim foot model with the extrinsic muscles; Model 3 (right) is the same as Model 2 but includes all the intrinsic foot muscles. Credit: Scarton A, Guiotto A, Malaquias T, Spolaor F, Sinigaglia G, Cobelli C, Jonkers I, Sawacha Z. A methodological framework for detecting ulcers’ risk in diabetic foot subjects by combining gait analysis, a new musculoskeletal foot model and a foot finite element model. Gait Posture. 2018;60:279 85.

Building on this model, Scarton et al. presented a methodological framework that used muscle activation predictions from a musculoskeletal model of the foot to drive a FE model [41]. They used a version of the previously described 8-degree of freedom model adapted to include the lumbrical muscles of the foot (Fig. 24.7). The authors confirmed the importance of the including the intrinsic foot muscles in these types of models as their results found improved predictions of plantar pressures when these were included. An inverse kinematic-based analysis was performed by Kim and Kipp [42] to estimate muscle and ligament strains from kinematic data for an OpenSim foot model with multiple segments during vertical hopping. They reported that the number of segments used (2, 3, or 5) did have a significant impact on the results, with the 2-segment model consistently overestimating ankle range of motion and tissue strains. Prinold et al. developed a patient-specific musculoskeletal model in OpenSim to estimate ankle joint forces in patients with juvenile idiopathic arthritis [43]. In testing the sensitivity of their modeling pipeline, they found that their results were highly dependent on the definition of the ankle joint axes and Achilles tendon insertion point. The authors also attempted to split the foot into two segments, assigning the ground reaction force vector first to the hindfoot and then, once the center of pressure passed the MTPJ flexion/extension axis, wholly to the toe segment. The results were inconsistent between participants. The elastic energy stored in the plantar aponeurosis (PA) during hindfoot and non-hindfoot strike running patterns has been investigated using a multi-segment foot model developed in OpenSim that included talus, calcaneus, midfoot, and toe segments [44]. Foot strike changed the pre-strain on the PA and its function during early stance but not the overall energy storage.

24.3

Areas of future biomechanical research

Musculoskeletal modeling remains a highly active area of research, with new techniques and models being developed and published on a regular basis. While more studies using musculoskeletal simulations that use detailed foot models are being reported, there remains significant research that needs to be performed in this area. A true high impact clinical application has yet to be demonstrated for detailed foot and ankle models, and this will be necessary to move this approach out of the research realm. Both forward and inverse dynamic simulations with foot models have the potential to help us better understand and therefore optimize clinical interventions, perhaps in the context of planning surgical interventions or in foot orthotic design. Integrating FE and musculoskeletal modeling (multiscale modeling) is another area with significant potential. This approach can allow us to simulate biomechanical problems with accurate boundary conditions. In theory, using joint and muscle forces from the musculoskeletal models can be used as inputs for a more detailed FE model. This approach has been applied in several body parts [45]; however, limitations still exist and further development is still needed. Improving the accuracy of biomechanical data collected as inputs to many models remains an important goal. Techniques to assign shear loading more accurately to multiple foot segments would be beneficial, as would improving the accuracy of capturing intrinsic foot joint motion, whether this be through the use of imaging technologies or accounting for the effect of skin motion artifacts to provide more accurate input kinematics.

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J Biomech 2012;45(14):2476 80. [7] Ma Y, Jiang S, Mithraratne K, Wilson N, Yu Y, Zhang Y. The effect of musculoskeletal model scaling methods on ankle joint kinematics and muscle force prediction during gait for children with cerebral palsy and equinus gait. Comput Biol Med 2021;134:104436. [8] Hill AV. The heat of shortening and the dynamic constants of muscle. Proc R Soc Lond Ser B - Biol Sci 1938;126(843):136 95. [9] Anderson FC, Pandy MG. A dynamic optimization solution for vertical jumping in three dimensions. Comput Methods Biomech Biomed Engin 1999;2(3):201 31. [10] Jamwal PK, Hussain S, Tsoi YH, Ghayesh MH, Xie SQ. Musculoskeletal modelling of human ankle complex: estimation of ankle joint moments. Clin Biomech 2017;44:75 82. [11] Vaughan CL. Are joint torques the Holy Grail of human gait analysis? Hum Mov Sci 1996;15(3):423 43. [12] Peng Y, Wang Y, Wong DW-C, Chen TL-W, Zhang G, Tan Q, et al. Extrinsic foot muscle forces and joint contact forces in flexible flatfoot adult with foot orthosis: a parametric study of tibialis posterior muscle weakness. Gait Posture 2021;88:54 9. [13] Steineman BD, Quevedo Gonza´lez FJ, Sturnick DR, Deland JT, Demetracopoulos CA, Wright TM. Biomechanical evaluation of total ankle arthroplasty. Part I: joint loads during simulated level walking. J Orthop Res Publ Orthop Res Soc 2021;39(1):94 102. [14] Thelen DG, Anderson FC. Using computed muscle control to generate forward dynamic simulations of human walking from experimental data. J Biomech 2006;39(6):1107 15. [15] Fluit R, Andersen MS, Kolk S, Verdonschot N, Koopman HFJM. Prediction of ground reaction forces and moments during various activities of daily living. J Biomech 2014;47(10):2321 9. [16] Dorn TW, Wang JM, Hicks JL, Delp SL. Predictive simulation generates human adaptations during loaded and inclined walking. PLOS ONE 2015;10(4):e0121407. [17] Haralabidis N, Serrancolı´ G, Colyer S, Bezodis I, Salo A, Cazzola D. Three-dimensional data-tracking simulations of sprinting using a direct collocation optimal control approach. PeerJ 2021;9:e10975. [18] Anderson FC, Pandy MG. Dynamic optimization of human walking. J Biomech Eng 2001;123(5):381 90. [19] De Groote F, Falisse A. Perspective on musculoskeletal modelling and predictive simulations of human movement to assess the neuromechanics of gait. Proc R Soc B Biol Sci 2021;288(1946) 20202432. [20] Dorn TW, Lin Y-C, Pandy MG. Estimates of muscle function in human gait depend on how foot-ground contact is modelled. Comput Methods Biomech Biomed Eng 2012;15(6):657 68. [21] Lopes DS, Neptune RR, Ambro´sio JA, Silva MT. A superellipsoid-plane model for simulating foot-ground contact during human gait. Comput Methods Biomech Biomed Eng 2016;19(9):954 63. [22] Wilson C, King MA, Yeadon MR. Determination of subject-specific model parameters for visco-elastic elements. 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[32] Oosterwaal M, Carbes S, Telfer S, Woodburn J, Tørholm S, Al-Munajjed AA, et al. The Glasgow-Maastricht foot model, evaluation of a 26 segment kinematic model of the foot J Foot Ankle Res [Internet] 2016;9Jul 8 [cited 2021 Mar 11]. Available from: https://www.ncbi.nlm.nih. gov/pmc/articles/PMC4938906/. [33] Kim Y, Lee KM, Koo S. Joint moments and contact forces in the foot during walking. J Biomech 2018;74:79 85. [34] MacWilliams BA, Cowley M, Nicholson DE. Foot kinematics and kinetics during adolescent gait. Gait Posture 2003;17(3):214 24. [35] Perry JE, Hall JO, Davis BL. Simultaneous measurement of plantar pressure and shear forces in diabetic individuals. Gait Posture 2002;15 (1):101 7. [36] Yavuz M, Erdemir A, Botek G, Hirschman GB, Bardsley L, Davis BL. Peak plantar pressure and shear locations: relevance to diabetic patients. Diabetes Care 2007;30(10):2643 5. [37] Saraswat P, Andersen MS, MacWilliams BA. A musculoskeletal foot model for clinical gait analysis. J Biomech 2010;43(9):1645 52. [38] Oosterwaal M, Telfer S, Tørholm S, Carbes S, van Rhijn LW, Macduff R, et al. Generation of subject-specific, dynamic, multisegment ankle and foot models to improve orthotic design: a feasibility study. BMC Musculoskelet Disord 2011;12(1):256. [39] Al-Munajjed AA, Bischoff JE, Dharia MA, Telfer S, Woodburn J, Carbes S. Metatarsal loading during gait-a musculoskeletal analysis. J Biomech Eng 2016;138(3):034503 034503 6. [40] Malaquias TM, Silveira C, Aerts W, Groote FD, Dereymaeker G, Sloten JV, et al. Extended foot-ankle musculoskeletal models for application in movement analysis. Comput Methods Biomech Biomed Eng 2017;20(2):153 9. [41] Scarton A, Guiotto A, Malaquias T, Spolaor F, Sinigaglia G, Cobelli C, et al. A methodological framework for detecting ulcers’ risk in diabetic foot subjects by combining gait analysis, a new musculoskeletal foot model and a foot finite element model. Gait Posture 2018;60:279 85. [42] Kim H, Kipp K. Number of segments within musculoskeletal foot models influences ankle kinematics and strains of ligaments and muscles. J Orthop Res 2019;37(10):2231 40. [43] Prinold JAI, Mazza` C, Di Marco R, Hannah I, Malattia C, Magni-Manzoni S, et al. A patient-specific foot model for the estimate of ankle joint forces in patients with juvenile idiopathic arthritis. Ann Biomed Eng 2016;44:247 57. [44] Wager JC, Challis JH. Elastic energy within the human plantar aponeurosis contributes to arch shortening during the push-off phase of running. J Biomech 2016;49(5):704 9. [45] Martelli S, Kersh ME, Schache AG, Pandy MG. Strain energy in the femoral neck during exercise. J Biomech 2014;47(8):1784 91.

Chapter 25

Predicting and Preventing Posttraumatic Osteoarthritis of the Ankle Donald D. Anderson1, Jason Wilken2, Claire Brockett3 and Anthony Redmond4 1

Departments of Orthopedics & Rehabilitation, Biomedical Engineering, and Industrial & Systems Engineering, University of Iowa, Iowa City, IA,

United States, 2Department of Physical Therapy and Rehabilitation Science, University of Iowa, Iowa City, IA, United States, 3Institute of Medical & Biological Engineering, University of Leeds, Leeds, United Kingdom, 4Clinical Biomechanics and Physical Medicine, University of Leeds, Leeds, United Kingdom

Abstract The vast majority of ankle osteoarthritis (OA) is the sequela of a joint injury; it is therefore by definition posttraumatic OA (PTOA). In extreme cases, the precipitating trauma may be an intra-articular fracture of the ankle. The trauma may be more subtle in other cases, such as damage to stabilizing ligaments leading to chronic ankle instability. Abnormal harmful biomechanics, or pathomechanics, explains much of the PTOA risk following a joint injury of the ankle. These pathomechanics arise from joint injury-associated acute mechanical overload, chronic cumulative mechanical overload, and/or altered joint kinematics. This chapter addresses these different pathomechanical states that challenge the ankle joint and can lead to PTOA. It also offers insight into a range of potential ways in which knowledge of the pathomechanical risk can be used to mitigate or prevent PTOA of the ankle.

25.1

Introduction: pathomechanical origins of posttraumatic osteoarthritis

Osteoarthritis (OA) is a painful and disabling condition involving the degeneration of the articular joint surfaces. In the ankle, the vast majority of OA occurs with a prior history of trauma to the joint. Such cases are defined as posttraumatic OA (PTOA). In more extreme cases, the precipitating joint injury can be that of an intra-articular fracture involving the ankle joint. This will most often involve fracture of the distal tibial plafond. The injury may be much more subtle in other cases, such as damage to the stabilizing ligaments of the ankle joint that lead to chronic ankle instability. There is a growing appreciation that abnormal harmful biomechanics, which is termed as pathomechanics here, explains much of the PTOA risk following a joint injury involving the ankle [1]. These pathomechanics may stem directly from the acute mechanical overload associated with the injury, from chronic cumulative mechanical overload that persists after a joint injury, or from altered joint kinematics in the wake of the injury. This chapter addresses these different pathomechanical states that challenge the ankle joint and can lead to PTOA. It also offers insight into a range of potential ways in which this knowledge of the pathomechanical risk can be used to mitigate or prevent PTOA of the ankle.

25.2

Pathomechanics I: acute joint injury severity

Intra-articular fracture of the tibial plafond is arguably the single type of joint injury that most likely leads to PTOA in the ankle [2]. PTOA rates after tibial plafond fracture have been reported to exceed 50% in several studies [3]. The hallmark signs of OA—joint space narrowing, sclerosis of subchondral bone, and the formation of osteophytes—are often seen within the first two years following these injuries [4]. This is despite surgical fracture reduction aimed at restoring Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00036-6 © 2023 Elsevier Inc. All rights reserved.

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normal joint biomechanics and function. The etiology of PTOA in these cases remains unclear, but pathomechanical factors are clearly implicated. Acute fracture severity is undoubtedly one of the primary factors in determining PTOA risk, but until recently, its objective quantification has been impossible. Instead, subjective fracture classification schemes, loosely based on severity but also including empirical anecdotal prognostic factors, are the means whereby the injuries are categorized. Fracture mechanics is a mature branch of physics that explains the relationship between the energy stored in a structure when loaded and the propensity for fracture. Failure initiates when the loading exceeds the capacity of a structure, and energy is then liberated in the fracturing process (Fig. 25.1). In the case of brittle solids (such as bone at high rates of loading), the energy absorption is proportional to the fracture-liberated surface area in the structure [5]. This provides a basis for objectively characterizing severity based on a physically grounded measure: fracture energy. The fracture-liberated surface area can be computed from clinical CT scan data that are routinely obtained in the course of treating these fractures [6 8]. The Hounsfield Unit values that are provided by CT also allow the surface area to be weighted by a bone density-based energy release rate to account for variation in bone quality (Fig. 25.2). The degree of articular comminution is a second unique fracture-related characteristic that can also be derived from these CT data. It can be defined based either on the percentage of fracture energy within some small distance of the articular surface [8] or on the length of all cracks in the articular surface arising from fracture [9]. This latter measure, the articular fracture edge length, holds attraction as a measure that builds upon ex vivo experimental testing that has shown preferential cell death along the fracture edge that later progresses across the articular surface [10]. Based on these fracture mechanics principles, objective techniques to measure the fracture severity from CT scan data were developed and validated first in surrogate bone specimens [11,12] and then in bovine bone specimens [6].

FIGURE 25.1 An intra-articular fracture of the tibial plafond occurs when an axial impact is sustained that exceeds the shock absorbing capacity of the bone. A CT volumetric rendering of a fractured tibia is shown to the right, with its intact contralateral tibia (flipped) to the left. Inter-fragmentary bone surfaces (red painted surfaces) are liberated in the fracture event (computer model of the fractured tibia shown in exploded view, center).

FIGURE 25.2 The area of inter-fragmentary bone surfaces is measured from CT using semi-automated methods (fragments reflected to visualize fracture surface, left). The fracture-liberated surface area and the bone densities across that surface are used to calculate fracture energy. The length of the edge between the subchondral and interfragmentary bone surfaces (the articular fracture edge length—highlighted with dashed black lines) can be used to quantify articular surface involvement. Original fragment pose from clinical CT scan, right.

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FIGURE 25.3 The rank orderings of severity based on the objective fracture energy-based metric reasonably agreed with subjective clinical rank orderings of severity. However, note the wide dispersion of rankings toward the lower end of the spectrum of severity. The fractures ranked 3 and 7 by severity objectively were ranked as much more severe by clinical experts. Neither of these patients later developed PTOA, while those whose fractures were ranked 10 and 13 objectively, but ranked much lower by the experts, both later developed PTOA.

Following this developmental work, the fracture energy methods were used to investigate the relationship between fracture severity and PTOA risk in an initial single-center patient series that involved 20 patients with tibial plafond fractures [7,8]. The severity of these fractures was also subjectively rank-ordered based on radiographs by three experienced orthopedic surgeons. The clinical outcomes for these patients were then collected at two years after surgical intervention. Kellgren-Lawrence (KL) scores [13] were used to characterize the radiographic OA status and the ankle OA scale score was obtained using a patient-reported questionnaire. First, reasonable agreement (74% concordance) was established between the subjective clinical rank orderings and the objective fracture energy-based measure (Fig. 25.3) [7]. There were a few notable exceptions where the subjective assessment of severity did not align well with the fracture energy-based measure. Interestingly, the fracture energybased measure gave much better prediction of KL scores (88% concordance) than did the clinical rank ordering (72% concordance—Fig. 25.4) [8]. In fact, when the PTOA status was converted to a binary measure (KL , 2, no PTOA and KL $ 2, PTOA present), there was a fracture energy-based severity threshold that was found to perfectly predict PTOA status. This was in contrast to 79% 84% PTOA prediction accuracy based on the rank orderings. Research work continues in this area, with full recognition that more cases need to be analyzed before finalizing this PTOA risk assessment tool, but these results strongly suggest that the acute fracture severity significantly influences PTOA risk after an intra-articular fracture involving the ankle joint. The surgical treatment of these injuries has long been focused on restoring articular congruity, and years of empirical clinical experience has been distilled into a dogma that emphasizes precise reduction to lessen the risk of PTOA [2]. The findings of concurrent studies conducted to understand the influence of contact stress aberration expected to accompany imprecise reduction upon PTOA risk are described in the following section.

25.3

Pathomechanics II: chronic stress aberration

The clinical premise that precise articular reduction can lessen the likelihood of PTOA is based on several studies involving patients with fractures of the tibial plafond. This has led surgeons to prioritize near-perfect reduction of the articular surface. In the ankle, with limited overlying soft tissues that were also injured, this pursuit of perfection translates into more time in the operating room and additional tissue handling, introducing substantial additional risk for complications [4]. Soft tissue complications, including infection, can lead very quickly to a need for amputation in the most severely injured extremities. Thus, there is a great impetus to better understand the pathomechanical origin of PTOA risk in these cases so that these varied surgical considerations can be factored into treatment decision-making. Prior experimental work involving cadaveric preparations established that contact stress elevations arise near incongruities of the articular surface [14,15]. These experimental constructs necessarily included highly idealized “step-off”

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FIGURE 25.4 (A) Clinical experts rank-ordered 20 tibial plafond fracture patients according to their assessed fracture severity, and the rankings were in 72% concordance with the KL grade of PTOA two years after the injury. (B) In contrast, there was an 88% concordance between an objective CT-based fracture severity metric and KL grade. The severity metric successfully discriminated between cases that developed PTOA and those that did not, in a threshold-like manner. Modified from Figure 13 on page 13 from Anderson DD, Marsh JL, Brown TD. The pathomechanical etiology of post-traumatic osteoarthritis following intra-articular fractures. Iowa Orthop J 2011;31:1 20.

configurations that poorly reflect the more complex incongruity seen in actual patients. For this reason, complementary clinical studies were conducted to investigate the relationship between malreduction, contact stress elevation, and PTOA risk using mathematical modeling approaches [16,17]. These studies took advantage of a series of radiographs collected in patients with developmental dysplasia of their hips who were followed clinically for decades. The concept of a cumulative contact stress overexposure arose from the observation that (1) not all contact stress is deleterious and (2) it is the chronic contact stress elevation acting over a period of time that leads to PTOA. Only when these concepts were both taken into consideration were investigators able to adequately predict PTOA risk attributable to the elevated contact stress. Modern tools for computational mechanical modeling that utilize postoperative CT data have enabled more comprehensive patient-specific investigation of these relationships. Key elements of this approach include the incorporation of patient-specific postoperative articular joint geometry and the simulation of full gait cycle loading in the ankle joint [18]. This has allowed the quantification of chronic aberrant contact stress exposure and its detrimental cumulative effect on PTOA risk [19].

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The basic premise of this work is that a contact stress-time exposure metric can be used to quantify the mechanical insult to the cartilage over time. Finite element analysis methods furthermore allow for the spatial distribution of contact stress (MPa) to be computed. Then local contact stress estimates can be evaluated further based on whether they exceed what is considered a reasonable damage threshold. The spatial distribution of per-gait-cycle cumulative contact stress-time overexposure dose (MPa-s) provides an index of the deleterious contact stress distributed over the articulating surface for a given discrete time within gait (Fig. 25.5). Utilizing finite element analysis methods and these concepts of contact stress-time overexposure, the same series of patients with tibial plafond fractures mentioned in the preceding section were studied [19]. The computed contact stress distributions (Fig. 25.6) clearly showed much larger percentages of the postfracture cartilage experienced high contact stress (areas above 7.5 MPa at the instant of peak joint loading averaged 23.4% of the surface in the fractured vs only 12.5% in the intact ankles). Furthermore, there was a clear distinction between area engagement histograms from those fractured ankles that developed PTOA within two years, versus those that did not, with the histograms of the latter being much more similar to those of intact contralateral ankles (Fig. 25.7). Putative contact stress damage threshold/tolerance parameters were systematically varied to determine the values that provided the best agreement between contact stress exposure metrics and the KL scores for each ankle. The predictive performance of five different contact stress exposure metrics was assessed in the intact versus fractured ankles, with fractured ankles grouped according to whether or not they developed PTOA by two-year follow up. The concordance

FIGURE 25.5 The contact stress-time exposure metric is computed as the spatial distribution of per-gait-cycle cumulative contact stress-time dose for pressures exceeding a damage threshold. This reflects the mechanical insult to the cartilage accumulating over time.

FIGURE 25.6 Inferior view of the contact stress exposure distribution on the tibial articulating surfaces for intact and fractured (reduced) ankles of the patients studied. Modified from Figure 4 on page 1042 from Li W, Anderson DD, Goldsworthy JK, Marsh JL. Brown TD. Patient-specific finite element analysis of chronic contact stress exposure after intraarticular fracture of the tibial plafond. J Orthop Res 2008;26(8):1039 1045.

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FIGURE 25.7 The percent of supra-threshold contact area exposed to given contact stress time dose levels shows a clear demarcation between fractured ankles that develop PTOA by 2-year follow-up versus all other ankles. The critical dose of stress time exposure (3 MPa-s) that provided the best concordance with outcomes is also indicated. Modified from Figure 4 on page 36 of Anderson DD, Van Hofwegen CJ, Marsh JL, Brown TD. Is elevated contact stress predictive of post-traumatic osteoarthritis for imprecisely reduced tibial plafond fractures? J Orthop Res 2011;29(1):33 39.

between the various contact stress exposure metrics and KL score were all excellent, exceeding 88% [19]. The concordance with OA status was even higher, with all metrics yielding greater than 94% agreement. The best concordance (KL score of 95%, OA status 100% prediction accuracy) was associated with a local stress-time exposure metric, which showed a clear separation between groups with different OA status attained with the local stress-time exposure metric. For some patients, the difference in exposure between intact and fractured ankles was relatively minor (fractured exposure within 40% of intact), while for others, the difference was dramatic (fractured exposure 900% of intact). Research work continues in this area, but these results strongly suggest that the chronic contact stress-time overexposure significantly adds to PTOA risk after an intra-articular fracture involving the ankle joint is surgically reduced. New discrete element analysis methods for computing patient-specific contact stress distributions have been developed for use in this area [20,21]. They represent an expeditious alternative to finite element analysis, with some of their inherent simplifications producing more robust convergence behavior, making them arguably better suited for use in studying much larger groups of patients. What these findings have not yet addressed is the role of any altered kinematics on the risk of PTOA in the ankle, whether after a fracture or after some other ligamentous disruption. The following section addresses these issues.

25.4

Pathomechanics III: altered kinematics

There are several soft tissue-related injuries and conditions that may result in altered kinematics of the ankle joint, which is associated with subsequent joint degeneration. Ankle sprain is one of the most common musculoskeletal injuries throughout the global population. Accounting for up to 75% of all ankle injuries, inversion sprains (lateral ankle sprains) are by far the most common mode of trauma to the ankle, occurring at a rate of 1 per 10,000 in the US on a daily basis [22]. In 30% 80% of lateral ankle sprains [23 25], patients continue to experience on-going impairment— whether decreased function or recurrent sprain. Whilst there are a range of factors to consider in defining chronic ankle instability (CAI), it is commonly believed that development of CAI is initiated by ankle ligament injury [26] and can lead to PTOA. Lateral ankle sprains typically occur due to excessive plantarflexion, internal rotation, and inversion of the hindfoot about an externally rotated lower limb soon after heel strike or initial foot contact with the ground [27]. These injuries occur with high energy and at high strain rates, with the lateral ankle ligaments typically experiencing strain of approximately 100%s21. Where the strain causes stress that exceeds the tensile strength of the ligament, damage may occur. These injuries are frequently associated with damage and rupture to supporting ligaments of the ankle, specifically the

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FIGURE 25.8 The proposed mechanism of PTOA after a severe ankle sprain. (A) During a typical lateral ankle sprain (inversion) the medial aspect of the talus likely impacts the tibial plafond, which may result in (B) a talar osteochondral lesion (OCL). Direct trauma to the articular surface can initiate progressive, irreversible joint destruction culminating in (C) late-state PTOA years to decades after the original injury. Reproduced with permission from Figure 1 on page 441 of Delco ML, Kennedy JG, Bonassar LJ, Fortier LA. Post-traumatic osteoarthritis of the ankle: a distinct clinical entity requiring new research approaches. J Orthop Res 2017;35(3):440 453.

anterior talofibular ligament, the calcaneofibular ligament, the posterior talofibular ligament (less common), and the deltoid ligaments. Lateral ankle sprains are considered to be the third most common cause of PTOA, with single sprain events reported in 13.7% of ankle PTOA cases [28]. During the sprain, it is postulated that the medial aspect of the talar dome could impact upon the inner surface of the medial malleolus, causing damage at the joint surface. This may result in the development of a medial osteochondral lesion (OCL), a disruption between the cartilage and underlying talar bone (Fig. 25.8) [29]. These may range from a small amount of local tissue bruising (with softening of localized cartilage) or increase in severity to an osteochondral fracture of the talus [30]. The lesions are associated with instability, inflammation and chronic pain, and result in a significant reduction in the patient’s quality of life [31,32]. The incidence of OCLs are likely to be masked as it is very difficult to diagnose damage through planar X-ray [33], and this medical imaging tool is most commonly used in emergency departments to assess for fracture and damage following ankle trauma. In addition, the symptoms associated with OCL (inflammation, chronic pain, and instability) can be attributed to other clinical issues related to the initial injury, therefore the real incidence of OCL has not accurately been determined. However, it has been reported that OCLs occur in approximately 6.8% of all ankle sprains [34], although one study has arthroscopically identified medial talar OCLs in 89% of cases with severe lateral ankle sprains [35]. Clearly, the changes in localized kinematics during a one-off sprain event can be sufficient to cause cartilage damage that contributes to long-term joint degeneration, with more than half of patients with OCLs developing OA. Although less frequent (accounting for 10% 20% of all sprains), tibiofibular syndesmosis sprains (otherwise called “high ankle sprains”) are more commonly reported in collision sport injuries and in sports where the ankle is restrained, such as skiing [36]. The mechanism for these injuries is typically external rotation of the ankle whilst the foot is in dorsiflexion [37]. The syndesmosis consists of four ligaments: the anterior inferior tibiofibular ligament, the posterior inferior tibiofibular ligament, the inferior transverse ligament, and interosseous ligament. They constrain motion of the fibula during gait to retain stability of the talus within the ankle mortise. Injury and disruption of these ligaments may increase the width of the ankle mortise, influencing the contact mechanics of the joint. Several cadaveric studies have explored this effect [38 40], generally demonstrating that under axial load alone, the natural bony geometry of the talocrural joint provides support that is sufficient to maintain joint stability. However, when the syndesmosis is disrupted, it can significantly change the contact mechanics of the joint. Then, once further motion was introduced, such as rotation of the joint, the contact area was seen to decrease significantly, and the contact pressure increase. Furthermore, a shift in the location of the center of pressure was also observed with syndesmosis injury, hence creating a combination of elevated stress and abnormal loading patterns within the ankle joint [40]—potentially leading to an adverse loading pathway as previously illustrated through ankle fracture. The incidence of one acute ankle sprain may predispose an individual to recurrent sprains, ongoing complications, and CAI. Saltzman et al. [28] identified that almost 15% of PTOA in the ankle resulted from recurrent ankle sprain. The reported prevalence of CAI following ankle sprain varies from 18% 70%—but certainly there appears to be a strong link [41,42]. There has been much debate on the paradigm of this condition and the pathology of ongoing instability, ranging from functional to mechanical impairment [43,44]. Contributions to CAI include: mechanical, neuromuscular, functional, and/or perceived deficits which can be presented long after acute sprain of the ankle. One of the more recent models for CAI [43] proposed three primary subgroups—mechanical instability, recurrent sprain, and neural

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perception of instability—that could exist either independently or coexist. Investigations of ankle motion, based on these groups, highlight that motion patterns are altered for mechanical instability, but not where individuals were in the perceived instability group alone [45]. Likewise, there is suggestion that mechanical instability may also be associated with anterior and inversion joint laxity [44]. Again, this change in joint stability is related to changes in joint kinematics—typically characterized by increased inversion at contact and a more inverted position of the ankle during running. A reduction in dorsiflexion has also been observed, with studies reporting malalignment of the hindfoot (anteriorly positioned talus with respect to a normal ankle), limiting the absorption of impact forces and increasing stress at the articular surface [46]. Similar to lateral ankle sprains, several studies have identified the evidence of articular cartilage damage associated with CAI. Magnetic resonance imaging techniques have been used through several studies and have highlighted a reduction in proteoglycan density and increased water content within talar cartilage, indicative of early pathogenesis [47]. However, they are insensitive to the identification of chondral lesions. OCLs of the talus have also been evidenced through arthroscopic measures taken during surgical procedure, with significant levels of degeneration observed in up to 95% of cases [35,48]. There appears to be a significant difference between the latency period of injury to end-stage PTOA when contrasting single sprain events and CAI, with single sprains showing signs of OA approximately 12 years sooner. It is proposed that the relatively rapid progression of degeneration following a single event injury may be related to the degree of cartilage damage sustained at the time of injury [28].

25.5

Areas of future biomechanical research

25.5.1 Posttraumatic ankle osteoarthritis: opportunities for intervention informed by pathomechanical knowledge Recent research findings suggest that early (nonsurgical) intervention could be beneficial in treating severe joint injuries in an effort to try and prevent PTOA [49 54]. Ongoing investigations are aimed at determining which interventions are the most promising therapeutic candidates. There are some indications that injecting an appropriate biological agent, such as amobarbital, which can target postinjury mitochondrial dysfunction, into the joint shortly following injury could be beneficial [51]. There are other indications that a mechanical intervention involving the wearing of a custom orthotic [55] or gait modification to limit aberrant joint kinematics [56] may be able to mitigate risk. Questions also arise regarding when and how to use a given intervention for greatest benefit, which joint injuries are best candidates for use in exploring these new therapeutic options, and how clinical trials can control for confounding pathomechanical factors. Moving forward, the evidence introduced in this chapter suggests that researchers should focus on determining how the pathomechanical concepts developed help to understand and explain PTOA risk. How else might this information be used? Could it be a basis for providing new tools for planning surgery? Could the information be used to provide intra-operative feedback to surgeons? Or could perhaps other precision medicine approaches be used to tailor treatments to the specific needs of a given injured joint? Metrics of acute fracture severity and elevated contact stress could both be useful in guiding treatment decisionmaking and advances in new treatments [50]. Building on successful animal work [51], clinical studies are beginning to be conducted to evaluate novel injectable therapeutic agents (e.g., amobarbital, dexamethasone) intended to mitigate or prevent PTOA [57]. It is likely that the success of the agent will vary depending on the severity and nature of the acute joint injury. This has implications for dosing, formulation, and the prescription of complementary agents that has begun to be studied with mathematical modeling [58]. Knowledge is also needed regarding how fracture severity influences modifiable targets in the critical degenerative pathway following intra-articular fracture. Additional work (detailed below) is aimed toward integrating computational modeling of ankle joint contact stress into the design and prescription of biological agents and ankle bracing that may reduce the contact stress and thereby mitigate PTOA risk [59]. Improving the precision of surgical fracture reduction could alleviate chronic contact stress elevation. An intraoperative biomechanical guidance system that builds on multifragment pose estimation from 2D fluoroscopic images has been developed (Fig. 25.9) [60]. The system uses well-established methods for pose estimation from 2D radiographic imaging and the tracking of 3D bone models [61,62]. In this specific application, the static pose of multiple bone fragments is derived from calibrated fluoroscopic images. This intra-operative tracking capability has most recently been evaluated in cadaveric experiments set in a mock operating room environment [63]. Five different intraarticular fracture patterns were produced in series of 5 cadaver ankles. Then each of the ankles was surgically reduced twice, once with the assistance of the biomechanical guidance system and once without. The two surgeries were

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FIGURE 25.9 Proposed intra-operative contact stress assessment system uses preoperative CT scans and intra-operative biplane imaging to determine the pose of fracture fragments, predict contact stress that would be associated with the fragments fixed in this pose, and display results to the surgeon. The surgeon can use these results to determine progress during the reduction.

separated by two days (to lessen the surgeon’s familiarity with the case), and the order in which the two surgeries were performed (guided vs unguided) was varied. The surgery involved closed reduction and percutaneous fixation of the fracture. Once the pose of the fracture fragments was obtained, computational contact stress analysis was performed to capture the degree of mechanical insult associated with the current state of reduction [21]. The results suggest that using a contact stress-based guidance system can allow surgeons to obtain a reduction that lessens the pathomechanical burden associated with residual malreduction of the articular surface. Surgical fracture reduction has been the mainstay of intra-articular fracture treatment for decades, but it is not the only way that joint contact stresses can be influenced. Furthermore, even with the best surgical effort, there often remains residual incongruity that leads to elevated contact stress [19]. Recent advances in bracing provide an attractive complementary treatment option. Custom dynamic orthoses (CDOs), such as the Intrepid Dynamic Exoskeletal Orthosis device design developed at Brooke Army Medical Center, have been used to dramatically improve function and reduce pain in hundreds of service members with traumatic limb injury [64 69]. The carbon fiber device is comprised of a proximal cuff just below the knee, a posterior strut used to store and return energy, a semirigid foot plate, and a heel cushion between the footplate and shoe [66]. These design components can be varied to influence the forces and motions experienced by the limb [70 77], which in turn influences the forces on the foot and activation of muscles that cross the ankle [71,72,74]. Recent findings indicate that this can decrease loading of the foot [78], suggesting that CDOs may be tuned to reduce articular contact stress (Fig. 25.10). Ongoing work aims to build upon a purpose-developed CDO design and manufacturing framework based on specific prescription models for patient-specific high-performance carbon fiber composite orthoses. The prescription model defines a CDO design approach, and patient-specific data are integrated into the prescription process. The software framework combines the prescription model with composite design, analysis, and manufacturing tools to generate customized carbon fiber composite CDO designs and manufacturing data. This framework is adaptable to various prescription models depending on CDO concept and requirements. Its utility has been demonstrated for two different prescription approaches and patient populations. The overall cycle time from prescription to patient fitting of the brace was less than 24 hours. The existing framework is now being modified to design and manufacture CDOs to prevent PTOA by decreasing harmful contact stress elevations at the ankle joint [59]. This involves developing a new PTOA-specific prescription model and identifying the patient-specific parameters that are of greatest importance. Objective measurement methods are being developed for these parameters to allow effective incorporation into the CDO design, engineering, and fabrication process. A key challenge for this effort is to derive engineering design requirements for a CDO that is suitable for PTOA patients, reduces contact stress, is customized using measured patient parameters, and effectively maintains or

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FIGURE 25.10 Vision of how custom dynamic orthosis design can be adapted to decrease harmful contact stress exposure.

restores mobility. Converting patient physical performance targets that mitigate or prevent PTOA to CDO engineering requirements will be achieved through transdisciplinary collaborations and preliminary simulations of CDO performance. It is expected that articular contact stress can be reduced through limb unloading ia the proximal cuff, reduction of the ankle joint moment and muscular forces through support provided by the posterior strut, and attenuation of impact and altered alignment from the heel wedge. With respect to limb unloading, forces on the limb commonly exceed 1.2 3 body weight while walking. These high forces spread over the small contact area of the ankle and contribute to high contact stresses. The external structure of a CDO allows forces applied at the proximal tibia to be transferred down through the posterior strut to the foot plate and ground, bypassing the ankle. Although partial limb unloading is a common goal in CDO fittings, the specific effects on articular contact stress are unknown. The proximal external support is being incorporated into an existing computational model to evaluate the effect of proximal cuff unloading on articular contact stress. Forces produced by muscles crossing the ankle are also important contributors to the total joint reaction forces and overall contact stress. Forces from the gastrocnemius and soleus muscles act at a distance from the joint to generate ankle joint moments, which control rotation of the lower limb and support and propel the body forward during gait [79]. The ankle joint contact forces resulting from muscle contraction are in addition to the forces from simple body weight support. As a result, the ankle joint experiences high joint reaction forces for much of the gait cycle. CDOs cross the ankle joint and can provide support, reducing the joint moment typically required by ankle musculature [80]. This reduction in moment could reduce muscle activation, the associated muscle produced joint reaction forces, and therefore joint contact stress. The computational model is being modified to replicate the ability of CDOs to reduce muscle activation and determine its effect on contact stress. Finally, joint alignment and the mechanical properties of the shoe-heel wedge-CDO interface (Fig. 25.10) can significantly influence device and limb loading [71,73]. Some effects are a direct response to the local mechanical change, while others are driven by the complex dynamics of walking and device patient interaction. Prior published findings are now being used to replicate the effects of limb orientation and loading in the existing ankle model. As part of pilot work in this area, the regional distribution of harmful contact stress exposures in the ankle fracture patients previously studied has been characterized by splitting their joint surfaces into 9 discrete and clinically relevant regions using a 3-by-3 grid from anterior to posterior and medial to lateral. These data were then queried to identify areas with increased contact stress exposure. The distribution of harmful contact stress exposures over the articular surface of the distal tibia was determined (Fig. 25.11). While the preponderance of elevated contact stress was located on the anterior third of the articular surface, additional sites of high exposure were identified centrally and mid-posteriorly.

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FIGURE 25.11 Plot showing regional variations in the amount of the distal tibia articular surface subjected to harmful contact stress time exposures as computed using finite element analysis. The colors of the columns are used to emphasize anterior third (blue), middle third (red), and posterior third (green) portions of the surface.

These findings suggest that the anterior joint surface, primarily loaded during late stance, is a good first target when seeking to reduce contact stress and the risk of PTOA development.

25.6

Summary/conclusion

PTOA is the most common form of ankle OA, which develops after an injury to the joint. Over the past decade, it has become increasingly clear that abnormal, harmful biomechanics, or pathomechanics, explains much of this PTOA risk. Pathomechanics take many forms, ranging from the sequela of subtle CAI to residual articular incongruity after an intra-articular fracture. As we begin to better understand these varied pathomechanical states in the ankle joint, we gain insight into a range of potential ways that we can mitigate or prevent PTOA of the ankle.

References [1] Anderson DD, Marsh JL, Brown TD. OREF Clinical Research Award: the pathomechanical etiology of post-traumatic osteoarthritis following intra-articular fractures. Iowa Orthop J 2011;31:1 20. [2] Dirschl DR, Marsh JL, Buckwalter JA, Gelberman R, Olson SA, Brown TD, et al. Articular fractures. J Am Acad Orthop Surg 2004;12(6): 416 23. [3] Zhang SB, Zhang YB, Wang SH, Zhang H, Liu P, Zhang W, et al. Clinical efficacy and safety of limited internal fixation combined with external fixation for Pilon fracture: a systematic review and meta-analysis. Chin J Traumatol 2017;20(2):94 8. [4] Marsh JL, Bonar S, Nepola JV, Decoster TA, Hurwitz SR. Use of an articulated external fixator for fractures of the tibial plafond. J Bone Jt Surg Am 1995;77(10):1498 509. [5] Von Rittinger PR. Lehrbuch der Aufbereitskunde. Berlin: Ernst and Korn; 1867. [6] Beardsley CL, Anderson DD, Marsh JL, Brown TD. Interfragmentary surface area as an index of comminution severity in cortical bone impact. J Orthop Res 2005;23(3):686 90. [7] Anderson DD, Mosqueda T, Thomas T, Hermanson EL, Brown TD, Marsh JL. Quantifying tibial plafond fracture severity: absorbed energy and fragment displacement agree with clinical rank ordering. J Orthop Res 2008;26(8):1046 52. [8] Thomas TP, Anderson DD, Mosqueda TV, Van Hofwegen CJ, Hillis SL, Marsh JL, et al. Objective CT-based metrics of articular fracture severity to assess risk for posttraumatic osteoarthritis. J Orthop Trauma 2010;24(12):764 9. [9] Dibbern K, Kempton LB, Higgins TF, Morshed S, McKinley TO, Marsh JL, et al. Fractures of the tibial plateau involve similar energies as the tibial pilon but greater articular surface involvement. J Orthop Res 2016;35(3):618 24. [10] Tochigi Y, Buckwalter JA, Martin JA, Hillis SL, Zhang P, Vaseenon T, et al. Distribution and progression of chondrocyte damage in a wholeorgan model of human ankle intra-articular fracture. J Bone Jt Surg Am 2011;93(6):533 9. [11] Beardsley CL, Bertsch CR, Marsh JL, Brown TD. Interfragmentary surface area as an index of comminution energy: proof of concept in a bone fracture surrogate. J Biomech 2002;35(3):331 8. [12] Beardsley CL, Heiner AD, Brandser EA, Marsh JL, Brown TD. High density polyetherurethane foam as a fragmentation and radiographic surrogate for cortical bone. Iowa Orthop J 2000;20:24 30. [13] Kellgren JH, Lawrence JS. Radiological assessment of osteo-arthrosis. Ann Rheum Dis 1957;16(4):494 502.

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[14] Baratz ME, Des Jardins Jd, Anderson DD, Imbriglia JE. Displaced intra-articular fractures of the distal radius: the effect of fracture displacement on contact stresses in a cadaver model. J Hand Surg [Am] 1996;21(2):183 8. [15] Brown TD, Anderson DD, Nepola JV, Singerman RJ, Pedersen DR, Brand RA. Contact stress aberrations following imprecise reduction of simple tibial plateau fractures. J Orthop Res 1988;6(6):851 62. [16] Hadley NA, Brown TD, Weinstein SL. The effects of contact pressure elevations and aseptic necrosis on the long-term outcome of congenital hip dislocation. J Orthop Res 1990;8(4):504 13. [17] Maxian TA, Brown TD, Weinstein SL. Chronic stress tolerance levels for human articular cartilage: two nonuniform contact models applied to long-term follow-up of CDH. J Biomech 1995;28(2):159 66. [18] Anderson DD, Goldsworthy JK, Shivanna K, Grosland NM, Pedersen DR, Thomas TP, et al. Intra-articular contact stress distributions at the ankle throughout stance phase-patient-specific finite element analysis as a metric of degeneration propensity. Biomech Model Mechanobiol 2006;5(2 3):82 9. [19] Anderson DD, Van Hofwegen CJ, Marsh JL, Brown TD. Is elevated contact stress predictive of post-traumatic osteoarthritis for imprecisely reduced tibial plafond fractures? J Orthop Res 2011;29(1):33 9. [20] Anderson DD, Iyer KS, Segal NA, Lynch JA, Brown TD. Implementation of discrete element analysis for subject-specific, population-wide investigations of habitual contact stress exposure. J Appl Biomech 2010;26(2):215 23. [21] Kern AM, Anderson DD. Expedited patient-specific assessment of contact stress exposure in the ankle joint following definitive articular fracture reduction. J Biomech 2015;48(12):3427 32. [22] Waterman BR, Owens BD, Davey S, Zacchilli MA, Belmont Jr PJ. The epidemiology of ankle sprains in the United States. JBJS 2010;92(13): 2279 84. [23] Hertel J. Functional anatomy, pathomechanics, and pathophysiology of lateral ankle instability. J Athl Train 2002;37(4):364. [24] Hiller CE, Nightingale EJ, Lin C-WC, Coughlan GF, Caulfield B, Delahunt E. Characteristics of people with recurrent ankle sprains: a systematic review with meta-analysis. Br J Sports Med 2011;45(8):660 72. [25] Agel J, Evans TA, Dick R, Putukian M, Marshall SW. Descriptive epidemiology of collegiate men’s soccer injuries: National Collegiate Athletic Association Injury Surveillance System, 1988 1989 through 2002 2003. J Athl Train 2007;42(2):270. [26] Doherty C, Bleakley C, Hertel J, Caulfield B, Ryan J, Delahunt E. Recovery from a first-time lateral ankle sprain and the predictors of chronic ankle instability: a prospective cohort analysis. Am J Sports Med 2016;44(4):995 1003. [27] Gehring D, Wissler S, Mornieux G, Gollhofer A. How to sprain your ankle a biomechanical case report of an inversion trauma. J Biomech 2013;46(1):175 8. [28] Saltzman CL, Salamon ML, Blanchard GM, Huff T, Hayes A, Buckwalter JA, et al. Epidemiology of ankle arthritis: report of a consecutive series of 639 patients from a tertiary orthopaedic center. Iowa Orthop J 2005;25:44. [29] Delco ML, Kennedy JG, Bonassar LJ, Fortier LA. Post-traumatic osteoarthritis of the ankle: a distinct clinical entity requiring new research approaches. J Orthop Res 2017;35(3):440 53. [30] O’Loughlin PF, Heyworth BE, Kennedy JG. Current concepts in the diagnosis and treatment of osteochondral lesions of the ankle. Am J Sports Med 2010;38(2):392 404. [31] Angthong C, Yoshimura I, Kanazawa K, Takeyama A, Hagio T, Ida T, et al. Critical three-dimensional factors affecting outcome in osteochondral lesion of the talus. Knee Surgery, Sports Traumatol Arthrosc 2013;21(6):1418 26. [32] Schachter AK, Chen AL, Reddy PD, Tejwani NC. Osteochondral lesions of the talus. JAAOS-J Am Acad Orthop Surg 2005;13(3):152 8. [33] Giannini S, Buda R, Faldini C, Vannini F, Bevoni R, Grandi G, et al. Surgical treatment of osteochondral lesions of the talus in young active patients. JBJS 2005;87(Suppl. 2):28 41. [34] Berndt AL, Harty M. Transchondral fractures (osteochondritis dissecans) of the talus. J Bone Jt Surg Am 1959;41(6):988 1020. [35] Taga I, Shino K, Inoue M, Nakata K, Maeda A. Articular cartilage lesions in ankles with lateral ligament injury: an arthroscopic study. Am J Sports Med 1993;21(1):120 7. [36] Amendola A, Williams G, Foster D. Evidence-based approach to treatment of acute traumatic syndesmosis (high ankle) sprains. Sports Med Arthrosc Rev 2006;14(4):232 6. [37] Williams GN, Jones MH, Amendola A. Syndesmotic ankle sprains in athletes. Am J Sports Med 2007;35(7):1197 207. [38] Beumer A, van Hemert WLW, Swierstra BA, Jasper LE, Belkoff SM. A biomechanical evaluation of clinical stress tests for syndesmotic ankle instability. Foot Ankle Int 2003;24(4):358 63. [39] Ogilvie-Harris D, Reed S, Hedman T. Disruption of the ankle syndesmosis: biomechanical study of the ligamentous restraints. Arthroscopy: J Arthroscopic Relat Surg 1994;10(5):558 60. [40] Hunt KJ, Goeb Y, Behn AW, Criswell B, Chou L. Ankle joint contact loads and displacement with progressive syndesmotic injury. Foot Ankle Int 2015;36(9):1095 103. [41] Gribble PA, Bleakley CM, Caulfield BM, Docherty CL, Fourchet F, Fong DT-P, et al. 2016 consensus statement of the International Ankle Consortium: prevalence, impact and long-term consequences of lateral ankle sprains. Br J Sports Med 2016;50(24):1493 5. [42] Hertel J. Functional instability following lateral ankle sprain. Sports Med 2000;29(5):361 71. [43] Hiller CE, Kilbreath SL, Refshauge KM. Chronic ankle instability: evolution of the model. J Athl Train 2011;46(2):133 41. [44] Hubbard TJ, Kramer LC, Denegar CR, Hertel J. Correlations among multiple measures of functional and mechanical instability in subjects with chronic ankle instability. J Athl Train 2007;42(3):361.

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[45] Brown C, Padua D, Marshall SW, Guskiewicz K. Individuals with mechanical ankle instability exhibit different motion patterns than those with functional ankle instability and ankle sprain copers. Clin Biomech 2008;23(6):822 31. [46] Wikstrom EA, Hubbard TJ. Talar positional fault in persons with chronic ankle instability. Arch Phys Med Rehab 2010;91(8):1267 71. [47] Song K, Pietrosimone B, Wikstrom E. Lower proteoglycan density within the talar articular cartilage is associated with worse postural control in individuals with chronic ankle instability. Osteoarthr Cartil 2018;26:S364 5. [48] Hintermann B, Boss A, Scha¨fer D. Arthroscopic findings in patients with chronic ankle instability. Am J Sports Med 2002;30(3):402 9. [49] Anderson DD, Chubinskaya S, Guilak F, Martin JA, Oegema TR, Olson SA, et al. Post-traumatic osteoarthritis: improved understanding and opportunities for early intervention. J Orthop Res 2011;29(6):802 9. [50] Buckwalter JA, Anderson DD, Brown TD, Tochigi Y, Martin JA. The roles of mechanical stresses in the pathogenesis of osteoarthritis: implications for treatment of joint injuries. Cartilage 2013;4(4):286 94. [51] Coleman MC, Goetz JE, Brouillette MJ, Seol D, Willey MC, Petersen EB, et al. Targeting mitochondrial responses to intra-articular fracture to prevent posttraumatic osteoarthritis. Sci Transl Med 2018;10(427). [52] Diekman BO, Wu CL, Louer CR, Furman BD, Huebner JL, Kraus VB, et al. Intra-articular delivery of purified mesenchymal stem cells from C57BL/6 or MRL/MpJ superhealer mice prevents posttraumatic arthritis. Cell Transpl 2013;22(8):1395 408. [53] Furman BD, Mangiapani DS, Zeitler E, Bailey KN, Horne PH, Huebner JL, et al. Targeting pro-inflammatory cytokines following joint injury: acute intra-articular inhibition of interleukin-1 following knee injury prevents post-traumatic arthritis. Arthritis Res Ther 2014;16(3):R134. [54] Olson SA, Furman BD, Kraus VB, Huebner JL, Guilak F. Therapeutic opportunities to prevent post-traumatic arthritis: lessons from the natural history of arthritis after articular fracture. J Orthop Res 2015;33(9):1266 77. [55] Quacinella M, Bernstein E, Mazzone B, Wyatt M, Kuhn KM. Do spatiotemporal gait parameters improve after pilon fracture in patients who use the intrepid dynamic exoskeletal orthosis? Clin Orthop Relat Res 2019;477(4):838 47. [56] Vincent KR, Conrad BP, Fregly BJ, Vincent HK. The pathophysiology of osteoarthritis: a mechanical perspective on the knee joint. PM R 2012;4(5 Suppl.):S3 9. [57] Grodzinsky AJ, Wang Y, Kakar S, Vrahas MS, Evans CH. Intra-articular dexamethasone to inhibit the development of post-traumatic osteoarthritis. J Orthop Res 2017;35(3):406 11. [58] Ayati BP, Kapitanov GI, Coleman MC, Anderson DD, Martin JA. Mathematics as a conduit for translational research in post-traumatic osteoarthritis. J Orthop Res 2017;35(3):566 72. [59] Anderson DD, Tanner BD, Wilken JM, Integrating pathomechanical risk of post-traumatic OA into the treatment of intra-articular fractures. In: XXVII Congress of the ISB, held in conjunction with the 43rd Annual Meeting of the ASB. Calgary, Canada; 2019. [60] Tatum M, Kern AM, Goetz JE, Thomas GW, Anderson DD. A novel system for near real-time markerless intra-operative bone tracking. Comput Methods Biomech Biomed Engin 2022. Submitted for publication. [61] Gong RH, Stewart J, Abolmaesumi P. Multiple-object 2-D-3-D registration for noninvasive pose identification of fracture fragments. IEEE Trans Biomed Eng 2011;58(6):1592 601. [62] Otake Y, Armand M, Armiger RS, Kutzer MD, Basafa E, Kazanzides P, et al. Intraoperative image-based multiview 2D/3D registration for imageguided orthopaedic surgery: incorporation of fiducial-based C-arm tracking and GPU-acceleration. IEEE Trans Med Imaging 2012;31(4):948 62. [63] Willey MC, Kern AM, Goetz JE, Marsh JL, Anderson DD. Biomechanical guidance can improve accuracy of reduction for intra-articular tibia plafond fractures and reduce joint contact stress. J Orthop Res 2022. Available from: https://doi.org/10.1002/jor.25393. In press. [64] Bedigrew KM, Patzkowski JC, Wilken JM, Owens JG, Blanck RV, Stinner DJ, et al. Can an integrated orthotic and rehabilitation program decrease pain and improve function after lower extremity trauma? Clin Orthop Relat Res 2014;472(10):3017 25. [65] Hsu JR, Owens JG, DeSanto J, Fergason JR, Kuhn KM, Potter BK, et al. Metrc, patient response to an integrated orthotic and rehabilitation initiative for traumatic injuries: the PRIORITI-MTF study. J Orthop Trauma 2017;31(Suppl. 1):S56 62. [66] Patzkowski JC, Blanck RV, Owens JG, Wilken JM, Blair JA, Hsu JR. Can an ankle-foot orthosis change hearts and minds? J Surg Orthop Adv 2011;20(1):8 18. [67] Patzkowski JC, Blanck RV, Owens JG, Wilken JM, Kirk KL, Wenke JC, et al. C. skeletal trauma research, comparative effect of orthosis design on functional performance. J Bone Jt Surg Am 2012;94(6):507 15. [68] Sheean AJ, Tennent DJ, Owens JG, Wilken JM, Hsu JR, Stinner DJ, et al. Effect of custom orthosis and rehabilitation program on outcomes following ankle and subtalar fusions. Foot Ankle Int 2016;37(11):1205 10. [69] Ikeda AJ, Fergason JR, Wilken JM. Clinical outcomes with the intrepid dynamic exoskeletal orthosis: a retrospective analysis. Mil Med 2019;. [70] Aldridge Whitehead JM, Russell Esposito E, Wilken JM. Stair ascent and descent biomechanical adaptations while using a custom ankle-foot orthosis. J Biomech 2016;49(13):2899 908. [71] Brown SE, Russell Esposito E, Wilken JM. The effect of ankle foot orthosis alignment on walking in individuals treated for traumatic lower extremity injuries. J Biomech 2017;61:51 7. [72] Harper NG, Esposito ER, Wilken JM, Neptune RR. The influence of ankle-foot orthosis stiffness on walking performance in individuals with lower-limb impairments. Clin Biomech (Bristol, Avon) 2014;29(8):877 84. [73] Ikeda AJ, Fergason JR, Wilken JM. Effects of altering heel wedge properties on gait with the intrepid dynamic exoskeletal orthosis. Prosthet Orthot Int 2017;42(3):265 74. [74] Ranz EC, Russell Esposito E, Wilken JM, Neptune RR. The influence of passive-dynamic ankle-foot orthosis bending axis location on gait performance in individuals with lower-limb impairments. Clin Biomech (Bristol, Avon) 2016;37:13 21.

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[75] Russell Esposito E, Blanck RV, Harper NG, Hsu JR, Wilken JM. How does ankle-foot orthosis stiffness affect gait in patients with lower limb salvage? Clin Orthop Relat Res 2014;472(10):3026 35. [76] Russell Esposito E, Choi HS, Owens JG, Blanck RV, Wilken JM. Biomechanical response to ankle-foot orthosis stiffness during running. Clin Biomech (Bristol, Avon) 2015;30(10):1125 32. [77] Russell Esposito E, Ranz EC, Schmidtbauer KA, Neptune RR, Wilken JM. Ankle-foot orthosis bending axis influences running mechanics. Gait Posture 2017;56:147 52. [78] Stewart J, Djafar T, Miltenberger R, Wyatt M. Plantar pressure changes with use of a custom dynamic ankle orthosis. In: 41st Annual Meeting of the American Society of Biomechanics. Boulder, CO; 2017. p. 282 283. [79] Neptune RR, Kautz SA, Zajac FE. Contributions of the individual ankle plantar flexors to support, forward progression and swing initiation during walking. J Biomech 2001;34(11):1387 98. [80] Arch ES, Stanhope SJ, Higginson JS. Passive-dynamic ankle-foot orthosis replicates soleus but not gastrocnemius muscle function during stance in gait: insights for orthosis prescription. Prosthet Orthot Int 2016;40(5):606 16.

Chapter 26

Mechanics of Biological Tissues Arturo Nicola Natali1,2, Emanuele Luigi Carniel1,2 and Chiara Giulia Fontanella1,2 1

Department of Industrial Engineering, University of Padova, Padova, Italy, 2Centre for Mechanics of Biological Materials, University of Padova,

Padova, Italy

Abstract The investigation of the mechanics of biological tissues is presented with regard to the biomechanical response of the foot in healthy and degraded conditions, as integrated with experimental and numerical approaches. A complete three-dimensional numerical model of the foot is reported considering bones, ligaments, adipose tissues, and skin. Constitutive models and associated parameters are defined to interpret the mechanical response of different biological tissues. The analyses consider typical features such as large displacements and strains, nonlinear elasticity, and time-dependent phenomena. The constitutive models are formulated starting from the analysis of the complex structural and micro-structural configuration of the tissues and evaluating the relationship between tissue histology and mechanical properties. The evaluation of the constitutive parameters is performed by a coupled deterministic and stochastic optimization method, considering data from experimental tests. The comparison of numerical model results and experimental data confirm the model’s ability to describe the mechanical behavior of the tissues investigated. Numerical analyses are performed to evaluate the biological tissues and structural mechanical behavior, while considering loading during the phases of the gait cycle. Moreover, degraded phenomena are investigated, considering the effects on the mechanical response. This chapter provides an overview of the methodology that integrates experimental and numerical approaches with biological tissues and structural mechanics, for a better interpretation of foot biomechanics, while supporting clinical activities.

26.1

Introduction

The human foot and ankle complex provides a stable base for standing, dampens the applied loads, and adapts to uneven ground in barefoot and shod conditions. This complex is composed of bones, cartilage, muscles, tendons, skin, and ligaments that operate during standing and locomotion. The biomechanical behavior of the foot can be investigated using combined experimental and computational approaches, developed considering the structural conformation of the foot and the proper mechanical response of the biological tissues. For the purposes of this chapter, the discussion will be limited to foot tissues, aiming to report on the procedures developed and to outline a methodology for foot biomechanical analyses. Computer modeling of the foot has been developed over the years to interpret the mechanical functionality of foot tissues and structures, to explore stress-related injuries, as well as to improve orthoses and footwear design [1
8]. As with other tissues of the musculoskeletal system, foot tissues can be affected by various degenerative factors including trauma, congenital, or inflammatory/infectious disorders that influence the mechanical response. For example, histomorphological alterations to the plantar fat pad induce significant modification of the tissue’s mechanical properties, and can lead to stiffening, lower strain threshold for injury, and lower damping capabilities that cause it to be unable to correctly perform its structural functionality [8
10]. Knowledge of the biomechanical properties of healthy and diseased feet may be used for screening patients (e.g., subjects with diabetes or heel pain subjects) and for the prevention of pathologies through the development of specific orthopedic devices [11,12]. Experimental approaches can be investigated and reciprocally integrated with a computational model of the foot. Such a model would allow a better understanding of the stress
strain relationship of the tissues, to evaluate phenomena that are not measurable with sufficient accuracy by means of experimental tests. The constitutive formulations are able to interpret the mechanical response of healthy feet and the influence of degenerative phenomena on tissue mechanics. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00023-8 © 2023 Elsevier Inc. All rights reserved.

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Computational modeling, if done correctly, is a validated and efficient method because of the ability to interpret many different situations while avoiding the technical and ethical difficulties of experimental testing. However, the development of reliable numerical models of biological structures is a very complicated and time-consuming activity. Historically, different simplifications in defining the geometry of the model have been made due to computational limitations. Moreover, the lack of tissue data, which are necessary for the definition of the tissue’s mechanical response, led to the choice of preliminary simplistic and approximate constitutive models. In the literature, two-dimensional models have been proposed, determining approximations in the evaluation of the overall biomechanical problem, with regard to the actual 3-dimensional motion or effective distribution of stress and strain [13
15]. Furthermore, simplified constitutive formulations have been assumed to characterize biological tissue mechanics, but proper formulations should be defined considering the actual conformation and functionality of the tissues. For example, three-dimensional models (3D) report a mechanical characterization of the adipose tissues by means of elastic or hyperelastic formulations [1,16,17], without considering the nonlinear viscoelastic behavior of the tissue [17
21]. Furthermore, numerical models of ankle ligaments reported in the literature [4,5,7] often accept linear elastic or isotropic hyperelastic characterization of tissue mechanical response and simplified geometrical representation of the structural morphometry, which can affect the model’s validity. The visco-hyperelasticity in the ligament tissue is an important aspect of their biomechanical response [22
27]. In the literature, several refined constitutive models are proposed for the mechanics of articular cartilage [28
30]. However, to preserve an accurate morphological representation and physiological loading, cartilage is often modeled as linear elastic, homogenous, and isotropic [6,31]. In this chapter, constitutive models are formulated starting from the analysis of the complex micro-structural configuration of the tissues, and then evaluating the relationship between tissue histology and mechanical properties. These models consider the typical features of the mechanical response of biological tissues, such as linear or nonlinear elasticity, time-dependent effects, geometrical nonlinearity, and anisotropy. A further topic of investigation pertains to constitutive parameter identification, which is performed by the inverse analysis of designed experimental tests. Finally, results from computational analyses are reported, aiming to show the potential of modeling techniques [14
20,31
33].

26.2

Materials and methods

In this section, different constitutive models are described to interpret the mechanical behavior of foot tissues. Specific attention is paid to the identification of the constitutive parameters that characterize the mechanical behavior of each tissue. Linear elastic, hyperelastic, and visco-hyperelastic constitutive formulations are usually assumed to interpret the mechanical behavior of different biological tissues. The identification of constitutive parameters is performed using a specific procedure based on the inverse analysis of experimental investigations.

26.2.1 Finite element modeling of the foot The definition, development, and validation of a numerical model of the foot entails a significant effort that requires the consideration of hard and soft tissues. Virtual geometric models of the bones are developed from computer tomography (CT) images, while geometric models of the soft tissue are mostly defined by the analysis of magnetic resonance (MR) images and histomorphometric data. The model considers the distal segments of the tibia and fibula, the full set of foot bones, the soft tissues which include cartilage, muscles, and connective tissues, as well as the plantar fascia, plantar soft tissue, and skin (Fig. 26.1) [14
20,31
33]. With regard to tarsal and metatarsal bony structures, the bones are often linked together by truss elements to mimic ligament function, while for more critical ligaments of the ankle joint complex, 3D representations of each ligament are more common (Fig. 26.2). The computational model is obtained by the finite element discretization of the solid model using a finite element pre-processor. Bone and soft tissues are meshed with tetrahedral elements that allow the morphometry of the system to be correctly interpreted, avoiding element distortion phenomena. The size of the elements ranged between 0.3 and 2.6 mm, with the smaller element in the regions where greater variability of stress and strain fields are expected.

26.2.2 Formulation of constitutive models The mechanical response of materials is defined by specifying the relationship between stress and strain, stress and strain rates, or other variables [34
37]. In the following, only isothermal problems will be considered. Some

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FIGURE 26.1 Representation of the numerical model of the foot with reference to bones, ligaments, adipose tissue, and planta fascia. The ligaments that join the distal epiphyses of tibia and fibula are the anterior and posterior tibiotalar ligaments (ATiFL and PTiFL). The deltoid region is composed by the anterior and posterior tibiotalar ligaments (ATTL and PTTL) and the tibiocalcaneal ligament (TCL), while the lateral-collateral region consists of the anterior and posterior talofibular ligaments (ATFL and PTFL) and the calcaneofibular ligament (CFL).

FIGURE 26.2 Solid model of components of the hindfoot region: (A) frontal and (B) posterior view of the ankle joint complex and (C) detail about the subtalar joint.

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constitutive schemes typically adopted in continuum mechanics are presented, together with the experimental observation of internal behavior, which lead to the constitutive model.

26.2.2.1 Linear elastic constitutive models Materials can show a mechanical response that is characterized by a linear relationship between stress and strain for infinitesimal strains [38
42]. This approximation is valid if body deformation is limited to an infinitesimal strain range and dissipative phenomena can be neglected. The most general relationship between stress and strain can be written as: σij 5 Dijkl εkl

ði; j; k; l 5 1; 2; 3Þ

(26.1)

where Dijkl are the 81 coefficients of the fourth order elasticity tensor, εij the components of the infinitesimal strain tensor and σij the elements of the Cauchy stress tensor. Not all 81 elastic constants Dijkl are independent since the elasticity tensor must fulfill some basic requirements. Because of the symmetry of strain and stress tensors, the following symmetries hold: Dijkl 5 Djikl 5 Dijlk

ði; j; k; l 5 1; 2; 3Þ

(26.2)

Further symmetries are due to the existence of a potential, which is the density of energy accumulated by a body because of its deformed state: @W Dijkl 5 (26.3) @εij @εkl These symmetry relationships are: Dijkl 5 Dklij

ði; j; k; l 5 1; 2; 3Þ

(26.4)

reducing the independent constants to 36. Any symmetry in the mechanical response of a material simplifies the form of the constitutive tensor. If a material shows the same response for every direction considered, it is said to be isotropic. In this case, it can be demonstrated that there are two independent constants and Eq. (26.1) can be rewritten as: 
  σij 5 μ δil δjk 1 δik δjl 1 λδij δkl εkl

ði; j; k; l 5 1; 2; 3Þ

(26.5)

where δij is the Kronecker delta and the terms μ and λ are Lame`’s elastic constants, related to the Young modulus E and to the Poisson ratio ν by the formulas: μ5

E 2ð 1 1 ν Þ

λ5

νE ð1 1 ν Þð1 2 2ν Þ

(26.6)

Isotropic models identify a very simple class of materials. This is not common in biological tissues, where materials tend to show a mechanical response related to their particular structural organization, depending on the orientation of some components. Two particular types of anisotropic materials are presented here: orthotropic materials and transversally isotropic materials. Orthotropic materials are characterized by a symmetry of behavior with respect to three orthogonal axes. This symmetry introduces some restrictions in constitutive relationships. 3 2 1 ν 21 ν 31 2 2 0 0 0 7 6 E1 E2 E3 7 6 7 6 ν 1 ν 32 6 2 12 2 0 0 0 7 7 6 E2 E3 72 3 6 E1 3 2 7 σ11 6 ε11 7 6 ν 13 ν 1 23 7 6 ε22 7 6 2 6 2 0 0 0 7 76 σ22 7 7 6 6 E3 E2 76 σ33 7 6 ε33 7 6 E1 ν ij ν ji 76 7 6 7 6 (26.7) 76 τ 12 7; with Ei 5 Ej ; i; j 5 1; 2; 3 6 ε12 7 5 6 1 76 7 6 0 7 6 0 0 0 0 74 τ 13 5 4 ε13 5 6 2G12 7 6 7 τ 23 6 ε23 1 7 6 6 0 0 0 0 0 7 7 6 2G13 7 6 6 1 7 5 4 0 0 0 0 0 2G23

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The elastic constants, Ei ; ν ij ; Gij are the Young moduli, the Poisson ratio, and the tangential moduli that specify the mechanical behavior along the principal directions, leading to nine independent elastic constants [43,44]. Transversally isotropic materials are characterized by the symmetry with respect to a specific direction, thus defining the plane of transversal symmetry. Under the hypothesis that direction 1 is normal to this plane of symmetry, the stress
strain relationship is simplified, reducing the number of independent elastic constants to 5: 8 E2 5 E3 > > > > ν > < 12 5 ν 13 ; ν 21 5 ν 31 G12 5 G13 ; G21 5 G31 > > E2 > > > : G23 5 1ð1 1 ν 23 Þ Transversal isotropy is a typical scheme for the mechanical characterization of long bones, such as the femur, tibia, etc. The thermodynamic consistency of constitutive laws restricts the values of the different elastic constants for orthotropic, transversally isotropic, and isotropic materials [38,39].

26.2.2.2 Hyperelastic constitutive models In the framework of the theory of hyperelasticity, a strain energy function must be defined to provide the stress
strain relationship [45,46]: SðCÞ 5 2

@W ðCÞ @C

(26.8)

where S is the second Piola-Kirchhoff stress tensor, C is the right Cauchy-Green strain tensor, and W is the Helmholtz free energy function. With regard to isotropic materials, the strain energy is function of the current strain state. Soft biological tissues usually contain a large liquid content. A direct consequence of this structural conformation is the fact that the tissue can behave like an almost-incompressible material. Therefore, a suitable computational framework requires the strain energy function and the stress response to be split into volumetric and volume-preserving (or isovolumetric) parts. The strain energy function is then defined by the following form [46]:

 Wm I~1 ; I3 5 Wmv ðI3 Þ 1 Wmi I~1

(26.9)

~ where I~1 is the principal invariant of the volume-preserving part of the right Cauchy-Green tensor C 5 I3 21=3 C, as ~ I~1 5 trðCÞ and I3 5 J 2 5 detðCÞ where J is the deformation Jacobian. Wmv is related to the volumetric part of strain and Wmi to the volume-preserving part. A typical formulation of the volumetric term that is able to account for the characteristics of soft biological tissues is reported [46]:  2 Kv I3 1=2 21 1 I3 2r=2 1 rI3 1=2 2 ðr 1 1Þ (26.10) Wmv ðI3 Þ 5 2 1 r ð r 1 1Þ where Kv is the bulk modulus in the unstrained configuration while r regulates volumetric compressibility with strain. The characteristic nonlinear response of the soft biological tissue, as outlined by experimental data, suggests the assumption of an exponential formulation for the volume-preserving term:
 C1 
  Wmi I~1 5 exp α1 I~1 2 3 2 1 α1

(26.11)

where C1 characterizes the shear stiffness of the tissue in the unstrained configuration, as G 5 2C1 , while α1 is a parameter that regulates the nonlinearity of the material response, with reference to experimental results. The contributions of fibers to the strain energy function must be defined. In the case of fibers aligned along one direction, transversally isotropic configuration can be assumed. The strain energy function can be reported as the following [47]:
 (26.12) W ðI1 ; I3 ; I4 Þ 5 Wmv ðI3 Þ 1 Wmi I~1 1 Wf ðI4 Þ

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The invariant I4 defines the anisotropy introduced by the fibers and specifies tissue stretch along the fibers’ direction as I4 5 λ2f . Fibers contribution can be described by considering their microstructural organization. In the unstrained configuration, fibers are usually characterized by a typical wavy conformation. When tensile load is applied, they first uncrimp and then get stretched. This mechanism determines a strongly nonlinear mechanical response that can be described by an exponential formulation of the fibers strain energy contribution [47
49]: Wf ðI4 Þ 5

 C4

exp½α4 ðI4 2 1Þ 2 α4 ðI4 2 1Þ 2 1 2 ðα4 Þ

(26.13)

where C4 is a constant that defines the initial stiffness of the fibers in the unstrained configuration, while α4 depends on the initial wavy conformation of fibers.

26.2.2.3 Visco-hyperelastic model Soft tissues exhibit a viscoelastic behavior that can be described by visco-hyperelastic models. The time-dependent behavior of soft biological tissues is due to the development of structural conformation rearrangements during loading. From a phenomenological point of view, typical examples of viscous effects are creep, stress-relaxation, and hysteresis. Rearrangement phenomena are usually defined as viscoelastic processes and can be associated with internal variables qi , which express material evolution during the stress
strain history from a phenomenological point of view. The mechanical state of the material is described by a specific configuration of the Helmholtz free energy function [46]:
 (26.14) ψ 5 ψ C; qi The specific formulation of the Helmholtz free energy can be developed accounting for mechanical models that are capable of describing the behavior of the material. Within viscoelastic theories the Zener model is frequently adopted. This model is made up of an equilibrium spring and viscoelastic branches connected in parallel. Every viscoelastic branch represents a viscoelastic process, which is characterized by a relative elastic stiffness γ i and a relaxation time τ i . The relative stiffness describes the contributions of the viscous processes to the whole instantaneous stiffness of the material, which is the stiffness during a straining process characterized by an infinite strain rate. The relative stiffnesses have to satisfy the following relationship [46]: n X γi 5 1 (26.15) γN 1 i51 N

where γ is the relative stiffness of the equilibrium spring, which defines the material behavior during an equilibrium strain process (i.e., a process that is characterized by a strain rate approaching zero). The analysis of the Zener model leads to the following formulation of the Helmholtz free energy [46,47]: n X

 ψ C; qi 5 W N ðCÞ 1 ψi C; qi

(26.16)

i51

where W N is an hyperelastic potential that defines the behavior of the equilibrium spring, while ψi is the Helmholtz free energy related to the ith viscous branch. In the present approach the variables qi describe the non-equilibrium stresses associated with the viscous processes and the following formulation of ψ_ i is assumed:
 1 i ψ_ i C; qi 5 W_ ðCÞ 2 qi :C_ 2

(26.17)

where W i is an hyperelastic potential associated with the ith spring, while the second term is the energy dissipated because of the viscous process. The potentials W N and W i can be related to an instantaneous hyperelastic strain energy W 0 , as W N 5 γ N W 0 and W i 5 γ i W 0 . The stress
strain relationship can be evaluated according to equation [46]: n  n X X

 @W N @W i 1 2 qi 5 SN ðCÞ 1 S C; qi 5 2 2 Si C; qi (26.18) @C @C i51 i51 The evolution law for viscous variables qi can be obtained by means of the Zener model: q_ i 1

1 i γ i @W 0 q 5 2 τi τ i @C

(26.19)

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26.2.3 Identification of constitutive parameters The definition of parameters that characterize the constitutive equations of biological materials is a challenge because of their complex mechanical relationships. The approach consists of an inverse analysis where the stress
strain history given by experimental procedures is obtained and parameter values that would define the best fit to the given data are estimated. Constitutive parameters are consequently evaluated using experimental data, corresponding model results, and optimization techniques. Mechanical tests should be performed on geometrically simple specimens and appropriate boundary conditions should be adopted to generate the most relevant stress
strain fields. Indeed, simple experimental tests can be described using analytical formulations. For more complicated situations, numerical methods must be adopted. Furthermore, experimental data should explore several different deformation modes to provide the necessary information for the characterization of the generic stress
strain behavior of the tissue and the identification of constitutive parameters. The procedure adopted for the definition of constitutive parameters requires the minimization of the discrepancy between experimental and analytical or numerical model results by means of a specific cost function. Several cost functions have been proposed in the literature. An example of a cost function that considers the weight of each data point in the output related to the ratio between the experimental data and model is [50]: vffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi u n  2 n
 1 uX Pii mod ðω;λi exp Þ Pii exp 1X 1 22 2 Θi ω;λi exp ; Pii exp ; Pii mod ΩðωÞ 5 t (26.20) exp exp mod n i51 Pii n i51 Pii ðω;λi Þ where ω is the set of constitutive parameters, n the number of experimental data points, λi exp the ith experimental input datum, Piiexp the ith experimental output value, and Piimod the ith analytical or numerical model output result corresponding to the constitutive parameters ω and the experimental input λi exp . The function ΩðωÞ is a measure of the overall discrepancy between experimental and model results when constitutive parameters ω are adopted considering thermomechanical restrictions on material behavior. Some limitations on constitutive parameters may be necessary. It may be difficult to define these conditions by boundaries of the domain parameters and should be more easily implemented by penalty contributions to the cost function, where the penalty term Θi assumes a variable value depending on the model result Piimod in relation to the criterion assumed. The optimization problem involves the evaluation of the set of constitutive parameters ωopt that minimizes ΩðωÞ. If the adopted constitutive model is strongly nonlinear, the cost function is characterized by multimodal behavior, and the function presents both a global minimum and further local minima. To get to a solution of the optimization problem by deterministic methods may result in the selection of local minima, without generating the optimal solution. The transition out of local minima is possible, from an operational point of view, by adopting stochastic algorithms, such as simulated annealing, making it possible to move toward the global minimum. Nevertheless, stochastic techniques do not guarantee to exactly reach local or global minima, but only to move close to minima themselves, which can be reached by introducing an additional deterministic step. The problem can be approached by coupled stochasticdeterministic optimization procedures, which allow for an efficient and reliable identification of the global minimum.

26.2.3.1 Constitutive parameter identification for bone The principal constituents of bone tissue are both organic (collagen fibers, non-collagenous proteins, and cells) and inorganic (hydroxyapatite crystals). Two principal types of bone tissue compose bony structures: cortical and trabecular. Cortical bone is a dense and compact mass, composed of cylindrical structures, known as osteons. The typical feature of trabecular bone is the reticular conformation, which is composed of a lattice of trabeculae. A cylindrical wrapping of bony lamellae builds up osteons and trabeculae. Lamellae are fiber-reinforced shells, with collagen fibers embedded within a ceramic matrix of hydroxyapatite. Each lamella contains fine fibers that run in approximately the same direction, but whose axes orientation can differ in adjacent lamellae. The microstructural configuration of bone tissues entails the anisotropic behavior, because of the orientation of osteons and trabeculae along principal loading directions. In this sense, the mechanical characterization of bone is usually performed using orthotropic linear elastic constitutive models [38,39], which require the identification of nine independent elastic parameters. To define average bony structures, parameter identification can be performed considering results from experiments that have been exhaustively developed on tissue specimens from different bony structures, considering both animal models and human samples [51]. With regards to patient specific models, refined frameworks have been provided to correlate CT data, that is, the bone volume fraction and the HU (Hounsfield Unit) values, and the orthotropic elastic parameters and directions [52,53].

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26.2.3.2 Constitutive parameter identification for cartilage Articular cartilage is a hierarchically organized tissue and has the important role to resist and distribute compressive and shear loads across joints, as well as providing a bearing surface with low friction. With concern to the histological aspect, articular cartilage is a specialized connective tissue composed of an extra-cellular matrix and a cellular part. The extra-cellular matrix is formed of a solid that includes proteoglycans and a fiber net of collagen and elastin. The fibers are embedded in interstitial fluid that is composed primarily of water (70%
85% of cartilage weight). The particular organization is optimized depending on the function of the tissue, resulting in four zones from the upper surface to the subchondral bone: tangential, medial, deep, and subchondral. On the tangential zone collagen fibers are organized parallel to the articular surface and radially distributed. This configuration is aimed at the proper response to compressive and shear stress [29]. The constitutive models proposed for its characterization are often very sophisticated, as multiphase models with a high number of parameters to be defined by means of experimental tests that prove to be very difficult to execute [29,30]. Considering the complexity of the anatomical site, it is necessary to choose an adequate constitutive model to have reliable numerical results, with a reasonable computational effort and a reliable definition of constitutive parameters, entailing the existence of adequate experimental reference data. The constitutive parameters of cartilage are determined from experimental data obtained from indentation tests [29,30]. The cartilaginous tissues of the model are characterized by the hyperelastic fiber-reinforced model reported in the previous section. According to the disposition of the collagen fibers, the unit vector is parallel to upper and lower surfaces and locally disposed in the radial direction [47]. A preliminary investigation of the standing position was developed applying a load of 765 N on the tibia axis (Fig. 26.3).

FIGURE 26.3 Schematic of an indentation test (A) and of the numerical disk sample with fiber direction (B). Results of hyperelastic fiberreinforced model compared with experimental data (C). Simulated displacement (D). Representation of stance condition minimum principal strain (E) and stress (F) in the hindfoot.

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26.2.3.3 Constitutive parameter identification for plantar soft tissue The plantar soft tissue is a complex system of connective and adipose components, characterized by a honeycomb configuration in which fibrous septa envelope adipose compartments [17]. The conformation of both fibrous and adipose components functions to maintain body stability, to attenuate ground impact, to redistribute plantar pressures, and to protect internal structures [17]. Plantar soft tissue mechanics can be interpreted by visco-hyperelastic formulations to account for nonlinear elasticity and time-dependent phenomena. The identification of parameters for healthy and degenerative conditions has been investigated previously [14,20]. The parameters were identified by the analysis of data from experimental investigations performed according to different loading conditions and sample configurations. A preliminary set of constitutive parameters has been evaluated using an optimization algorithm that minimizes the discrepancy between model results and experimental data from compression tests of cadaveric healthy tissue samples [17,46]. To interpret the mechanical response of living tissues, which is influenced by the interaction and connection with surrounding tissues and structures, constitutive parameters must be updated by considering data from indentation tests developed on the overall foot structure [46]. Because of the complex configuration of experimental tests, such analysis must be performed using numerical methods. Different sets of constitutive parameters can be defined starting from the

FIGURE 26.4 Sagittal plane X-ray image of a specimen in the clamp and view of the exposed fat pad (A). Histology of plantar soft tissues (B
C). Comparison of experimental and numerical results from indentation tests: (D) superposition of minimum discrepancy regions for low and high strain rate tests and identification of a set of parameters, results from (E) low and (F) high strain rate indentation tests. Models of the calcaneal region of a foot (G): contours of minimum principal stress for tests using a cylindrical indenter, developed at (H) low and (I) high strain rates.

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FIGURE 26.5 Results from in silico indentation tests (A). Numerical results (black continuous line) are reported within the domain of in vivo experimental data (gray region). Contours of the vertical displacement (B) and the minimum principal stress (C) fields are reported on a transversal section when a 5 mm indentation depth is applied.

preliminary set obtained, and adopted for the numerical analysis. By evaluating the discrepancy between numerical results and experimental data, it is possible to compute a domain of parameters where the discrepancy assumes minimal values, as a region of parameters that are capable of representing the mechanical response of the plantar soft tissue in healthy condition (Figs. 26.4 and 26.5) that also depends on strain rate. Aging or degenerative effects lead to histomorphological and mechanical changes in the plantar soft tissue. Degenerative phenomena can result in a gradual loss of collagen, a decrease in the quantity of elastic fibrous components, and a reduction of liquid content. Pathologies, such as diabetes, or atrophied phenomena, can cause alterations in saturated and unsaturated fatty acids. Some histological investigations show that the associated persistent hyperglycemia and the accelerated accumulation of advanced glycation end-products (AGEs) are associated with damage to fibrous tissue septa. The connective septa became thicker, contain a slightly higher percentage of fibrous tissues, and assume a fragmented configuration [9
12,54
58]. Experimental tests on foot tissues and structures in degenerative conditions confirm the presence of stiffer soft tissue, a less smooth plantar pressure distribution with higher local values, and a reduction in the ability to absorb shock phenomena [11,21,57]. To evaluate the influence of degenerative phenomena on the derived parameters, results from compression tests on degenerated tissue samples are analyzed [54
56], showing the increase of tissue initial stiffness and the reduction of the stress
strain nonlinearity, while viscous phenomena are not significantly affected. Once the qualitative trend of tissue mechanics with degenerative conditions is evaluated, the evolution of the parameters domain is identified by processing experimental results from in vitro tests. The procedure allows identification of a domain of constitutive parameters able to interpret the mechanical response of the tissue in specific degenerative conditions [45].

26.2.3.4 Constitutive parameter identification for ankle ligaments Ligamentous tissue is characterized by composite configuration, as collagen fiber bundles embedded within an isotropic ground substance matrix [47]. Transversally isotropic behavior is assumed because fibers are locally distributed along one main preferential direction. The conformation of collagen fibers and the interaction phenomena with ground substance entail nonlinear elastic and time dependent phenomena. The overall mechanical behavior is interpreted by a fiber-reinforced visco-hyperelastic constitutive model [47,49]. Several studies report data about the force-elongation behavior of ankle ligaments. Experimental specimens were prepared from fresh cadaveric human ankles to achieve the specific bone-ligament-bone structure [46,59
61]. In detail, soft tissues were removed to leave the tibia, fibula, talus, calcaneus, and their ligamentous attachments. Specimens were prepared to perform elongation tests on the anterior and posterior talofibular ligaments (ATFL and PTFL), the calcaneofibular ligament (CFL), the anterior and posterior tibiotalar ligaments (ATTL and PTTL), tibiocalcaneal ligament, and the anterior and posterior tibiotalar ligaments (ATiFL and PTiFL). Again, a multistep procedure can be assumed to identify a domain of parameters. A preliminary set of parameters is evaluated by assuming an analytical formulation to interpret experimental tests. Subsequently, computational methods are exploited to interpret the complexity of experimental results and fit different sets of parameters (Fig. 26.6) [49,61].

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FIGURE 26.6 Representation of the ankle ligaments with regard to the direction of the fibers by arrows: (A) solid model of the ankle structure, (B) detail of the conformation of the posterior talofibular ligament (PTFL), (C) fiber distribution in the PTFL solid models and in different ankle ligaments (D). Comparison between numerical results (continuous lines) and experimental data (open circles) of the lateral-collateral ankle ligaments: anterior talofibular ligament ATFL (E) and calcaneofibular ligament CFL (G). Contours of the displacement field at ATFL (F) and CFL (H).

26.2.4 Numerical analyses of foot functionality Computational analyses have been conducted, demonstrating the capabilities of the modeling techniques described above. An integrated experimental and numerical approach to biological tissues and structure mechanics was developed utilizing the general-purpose finite element software ABAQUS Standard and Explicit (Dassault Syste`mes Simulia Corp., Providence, RI). Constitutive models are implemented with the use of specific subroutines.

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26.2.4.1 Ankle movements Numerical analyses were performed to evaluate ankle response for different movements [49,61]. Dorsiflexion and plantarflexion were defined by fixing the calcaneus and imposing a rotation of the tibia and fibula around the intermalleolar axis in a range between 220 and 20 degrees. As reported in the literature, the intermalleolar axis is defined as the line that joins the medial and the lateral malleolus. Numerical analyses that interpret the inversion and eversion movements were also performed. In this case, the movements are defined fixing the calcaneus and imposing a rotation of the tibia and fibula around the longitudinal axis of the foot. This axis is defined as the line perpendicular to the plane containing the lateral malleolus, the medial malleolus, and the centroid of the tibial cross-section. The origin is located at the midpoint between the medial and the lateral malleolus. The numerical results allow identifying the roles of each ankle ligament during a range of ankle motions studied (Fig. 26.7). Ankle injuries are common in athletes, representing 25% of all sports-related injuries. 70% of ankle injuries are sprains and 40% lead to chronic symptoms [27]. Ankle sprains are correlated with ligament damage, in particular to the ATFL and CFL. The ATFL is the ligament that is most often damaged, while damage to both ATFL and CFL occurs in 20% of situations. The rupture of the CFL only or PTFL only is unusual. In the literature, several investigations of ankle injuries are reported, considering both experimental and computational techniques [62
66]. Computational analyses were performed considering the anterior drawer test and the talar tilt test, which are the common clinical examinations aimed at determining the integrity of the ankle ligaments [62,63]. The anterior and posterior motion along the longitudinal axis of the foot is defined as the displacement obtained by a load that ranges between 2125 and 125 N. The inversion and eversion rotation around the longitudinal axis of the foot was defined as the angular displacement produced in

FIGURE 26.7 Numerical results for a range of ankle motions: comparison between experimental and numerical results (A
B) and contours of the maximum principal strain for an anterior displacement of 8 mm reported on the frontal view of the ankle (C) and for an inversion of 17 degrees reported on the lateral view of the ankle (D). anterior talofibular ligament (ATFL) and anterior tibiotalar ligament (ATTL).

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response to a torque from 2 4 to 4 Nm. The results of numerical analyses were compared to experimental data in terms of force-vs-displacement and moment-vs-angle. The numerical results allow for identification of the roles of each ankle ligament during the ankle motions studied. Further numerical analyses of the anterior drawer and the inversion test after the rupture of some ankle ligaments were performed. In detail, the investigation entails the removal of the ATFL, ATTL, and CFL ligaments that reach the highest values of maximum principal strain, and considering results from experimental activities. With regard to the anterior drawer test, with a force of 125 N, the removal of ATFL results in an increase of anterior displacement of 13%. For the inversion-eversion test, the removing of ATFL and CFL results an angular rotation increase of 18% under a torque of 4 Nm [61].

26.2.4.2 Gait cycle Specific numerical analyses have been developed to evaluate the mechanical response of the foot and ankle tissues from heel strike to midstance [14]. In detail, the X-axis and the Z-axis were defined as two perpendicular axes on the plane of the ground. The perpendicular to that plane was the Y-axis. The center of the coordinate system was located in the center of the heel region on the skin. To evaluate the mechanical response of the foot during the initial phase up to heel strike, the sagittal-plane angle between foot and plane and the time of the analysis were defined considering results from the average experimental data. In detail, the transient peak of the heel strike was estimated to occur at 0.033 seconds of the gait cycle, and the angle between foot and plane was around 14 degrees. A relative motion between the foot and ground was imposed along Y direction. The friction between the skin and the rubber layer was modeled using contact surfaces with a friction coefficient of 0.61 as reported in the literature [14]. To interpret the phase between the heel.strike and the midstance, a rotation of the plane of 14 degrees around Z-axis was applied, while the tibia and the

FIGURE 26.8 Solid models of the foot and reference floor (A). Results from the numerical analysis: vertical force vs stance phase (B) Results from the numerical analysis: contours of the distributions of the minimum principal stress for the heel strike (C) and the midstance of the gait cycle (D), on a section at the second metatarsal.

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fibula were considered fully fixed. The time imposed for the motion was of 0.198 seconds, as shown by the experimental data. The minimum principal stress was reported at heel strike and midstance (Fig. 26.8). Higher stress values pertain to the plantar soft tissue, which was able to attenuate ground impact, redistribute plantar pressures, and protect internal structures. The results showed the increase of stress with the progression of the gait cycle. Numerical analysis were developed for the interpretation of the characteristic phenomenon of push off (Fig. 26.9) [67], showing the displacement field during the push off and the minimum principal stretch on a section of the third metatarsal head at subsequent phases. The results of the numerical analyses, in terms of stress and strain, make it possible to evaluate the mechanical role of the plantar soft tissue, distributing stress/deformation effects.

26.2.4.3 Foot and footwear interaction A refined finite element model of insole and sole of a referential walking shoe was developed (Fig. 26.10). The numerical model was developed to describe the real structural conformation and adopt adequate constitutive formulations for the different tissues and materials [67]. The numerical models of insole and sole are reported. Numerical analysis that interprets the mid-stance phase of the gait cycle was performed. Consequently, to simulate this condition, the upper surfaces of the tibia and fibula were fixed while a vertical displacement was applied to the ground support. The numerical analysis concludes when the reaction force evaluated on the tibia and fibula reached a load equivalent to 100% of body weight (750 N, 130 N on the fibula, and 620 N on the tibia) [68]. The foot-insole interface was modeled using contact surfaces with friction coefficients of 0.51, according to data in the literature [15]. The contours of the displacement and the compressive stretch within the soft tissues of the foot were determined (Fig. 26.11). The compressive stretch obtained on the upper surface of the insole was reported together with the minimum principal stress on the plantar region of the foot. The present analysis represents a general example of a method for investigating foot-insole interaction phenomena. The computational approach allows evaluating the performance of shoe components considering different stiffness for the insole or midsole, offering the possibility of influencing manufacturing procedures.

FIGURE 26.9 Results from the numerical analysis of the push off phase. The minimum principal stretch field is reported over a section at the third metatarsal head considering subsequent steps of the push off (A
C).

FIGURE 26.10 Representation of the solid model of the insole, midsole and outsole (A). Level of detail of the numerical model of the insole (B), midsole and outsole (C).

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FIGURE 26.11 Contour representation of the displacement magnitude (A) and compressive stretch evaluated along a longitudinal section of the foot (B) at the level of the second metatarsal. Contour representation of the compressive stretch evaluated on the upper surface of the insole (C) and contour of the minimum principal stress on the foot plantar region (D).

26.2.4.4 Diabetic condition Specific numerical analyses were developed to evaluate the influence of the disease progression on the overall mechanical response of the foot structure [45]. In particular, the analyses were performed considering the standing foot on rigid ground, when the force acting on each foot was around half of body weight. The interaction phenomena between ground and foot skin were specified by a Coulomb contact condition, according to a 0.42 friction coefficient. Numerical results were extrapolated assuming a specified deformation condition, to identify the influence of diabetes on the stress field (Fig. 26.12). Greater stress values pertained to greater degenerative conditions, showing the typical stiffening phenomenon.

26.2.5 Limitations of computational modeling The investigation of foot biomechanics using computational modeling must address many relevant problems related to the constitutive identification of tissue mechanics and the complex structural configuration. The constitutive modeling is rather complicated, but this complexity is essential for a proper interpretation of the mechanical behavior of the tissues. The proper identification of orthotropic elastic, hyperelastic, and visco-hyperelastic models used for representing foot tissues requires different experimental tests on human tissues which are not always available and at disposal. In the proposed numerical model, the relevance of the computational effort due to the effects of coupled material, geometric nonlinearities, and finite element discretization requirement must also be considered. A limitation of the proposed numerical model is that it corresponds to only one-foot morphometry, in terms of conformation and dimension. This condition is due to the relevant efforts required to develop specific models of different foot structures and tissues, composed of bones, ligaments, cartilage, soft tissues, and skin. Nevertheless, the specific conformation of foot structures has been defined and morphed accounting for average anthropometric data. The mechanical behavior of various foot tissues is identified considering experimental data developed on tissue samples from many different subjects. This action makes it possible to elaborate an average mechanical response of the soft tissues and consequently an average

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FIGURE 26.12 Domain of constitutive parameters for healthy and diseased conditions (A). Numerical results for different sets of constitutive parameters from healthy to diseased conditions (B). Results from the numerical analysis: contours of the magnitude displacement (C) and of the minimum principal stretch for incremental levels of degeneration (D
E), on a section at the second metatarsal level during the stance phase.

numerical model of the foot. In this context, model results do not interpret the mechanical behavior of the tissue from a specific subject foot, but the average mechanical response of foot tissues and structures.

26.2.6 Future biomechanics research Computational biomechanics provides structural modeling tools that allow for simulating: (1) the mechanical behavior of anatomical regions, (2) the interaction phenomena occurring among such biological structures, and (3) treatments, such as surgical procedures or prosthetic devices. The accuracy and reliability of computational methods are continuously improving because of the advancements in modeling research and the increase in computational power. Orthopedic devices deeply influence stress and strain distribution within foot. This variability can be interpreted by means of different constitutive models of materials, while geometric configuration can be varied within the numerical model. In this sense, the combination of experimental and computational approaches makes it possible to investigate the influence of orthopedic materials and devices on foot tissues mechanics. The analysis of the biomechanical behavior of the foot has a relevant socio-economic impact because of the increasing evidence of foot problems related to pathologies such as diabetes and obesity, and/or aging. In this way, a consideration must be given to the possibility of degenerative phenomena of hard and mostly soft tissues being accounted for within the constitutive formulations. In this sense, complex constitutive formulations, as visco-hyperelastic damage models, can be provided with regard to soft biological tissues, but more extended experimental investigations are required for the reliable application to foot biomechanics. While averaged models are mostly employed in the field of biomedical research, the focus of patient--specific models should be considered the future challenge. The principal tasks pertain to the validation of patientspecific computational models against in vivo results, the identification of model parameters, the solving time because of model complexity, the availability of powerful and efficient hardware and software, and the requirement of dedicated technicians and engineers.

26.3

Conclusion

In this chapter, we introduce the reader to the formulation adopted within numerical approaches to tissue mechanics and, in particular, to foot biomechanics. Attention is given to the definition of constitutive models, which can be used

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in the modeling of both hard and soft biological tissues. The reliability of numerical models is closely related to the affinity between the effective stress
strain response shown and the mathematical stress
strain relationships. Because biological tissues show complex behavior, with inelastic effects, such as viscoelasticity, the mathematical framework that must be adopted for an appropriate modeling must address this complexity. An additional task pertains to the fact that soft tissues undergo large strains. A combined experimental-numerical approach offers the possibility of a profitable integration of data, for both material characterization and result validation. In general, 3D numerical models allow for a possible representation of variable loading conditions, tissue conformation and properties, and offer a large set of results, in comparison with the repetitive and difficult action required by experimental testing. The accuracy of results in terms of strain and stress, that are often not easily accessible for experimental analysis, confirms the potential of the numerical approach that in any case requires the basic information from mechanical testing. The numerical results achieved offer an assessment of the overall procedure developed over several years. The present approach represents the basis for an extension in the evaluation of the biomechanical behavior of the overall foot components to the response in degenerative conditions such as diabetes. Moreover, the procedure can be associated with a computational tool for the investigation of the biomechanics of the healthy and degraded conditions in relation to the interaction phenomena occurring between foot and footwear products, to estimate the proper distribution of stress and strain inside the tissues.

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65. [39] Natali AN, Pavan PG, Carniel EL, Lucisano ME, Taglialavoro G. Anisotropic elasto-damage constitutive model for the biomechanical analysis of tendons. Med Eng Phy 2005;27:209
14. [40] Schwartz-Dabney CL, Dechow PC. Accuracy of elastic property measurement in mandibular cortical bone is improved by using cylindrical specimens. J Biomech Eng—T Asme 2002;124(6):714
23. [41] Schwartz-Dabney CL, Dechow PC. Variations in cortical material properties throughout the human dentate mandible. Am J Phys Anthropol 2003;120(3):252
77. [42] Schwartz-Dabney CL, Dechow PC. Edentulation alters material properties of cortical bone in the human mandible. J Dent Res 2002;81:613
17. [43] O’Mahony AM, Williams JL, Katz JO, Spencer P. Anisotropic elastic properties of cancellous bone from a human edentulous mandible. Clin Oral Implant Res 2000;11:415
21. [44] Natali AN, Carniel EL, Pavan PG. PG, Constitutive modeling of inelastic behaviour of cortical bone. Med Eng Phys 2008;30(5):905
1012. [45] Fontanella CG, Carniel EL, Natali AN. Numerical analysis of the foot in healthy and degenerative conditions. J Mech Med Biol 2017;17 (6):1750095. [46] Natali AN, Fontanella CG, Carniel EL. Constitutive formulation and numerical analysis of the heel pad region. Comput Methods Biomech Biomed Eng 2012;15(4):401
9. [47] Forestiero A, Carniel EL, Venturato C, Natali AN. Investigation of the biomechanical behaviour of hindfoot ligaments. Proc Inst Mech Eng H 2013;227(6):683
92. [48] Pavan PG, Stecco C, Darwish S, Natali AN, De R. Caro, Investigation of the mechanical properties of the plantar aponeurosis. Surg Radiol Anat 2011;33:905
11. [49] Forestiero A, Carniel EL, Natali AN. Biomechanical behaviour of ankle ligaments: constitutive formulation and numerical modeling. Comput Methods Biomech Biomed Eng 2014;17(4):395
404. [50] Natali AN, Forestiero A, Carniel EL. Parameters identification in constitutive models for soft tissues mechanics. Russian J Biomech 2009;4 (46):29
39. [51] Cowin SC. Bone mechanics handbook. 2nd ed. CRC Press; 2001. [52] Fritsch A, Hellmich C. ’Universal’ microstructural patterns in cortical and trabecular, extracellular and extravascular bone materials: micromechanics-based prediction of anisotropic elasticity. J Theor Biol 2007;244:597
620. [53] Toniolo I, Salmaso C, Bruno G, De Stefani A, Stefanini C, Gracco ALT, et al. Anisotropic computational modeling of bony structures from CT data: an almost automatic procedure. Comput Methods Prog Biomed 2020;189:105319. [54] Pai S, Ledoux WR. The compressive mechanical properties of diabetic and non-diabetic plantar soft tissue. J Biomech 2010;43:1754
60. [55] Pai S, Ledoux WR. The quasi-linear viscoelastic properties of diabetic and non-diabetic plantar soft tissue. Ann Biomed Eng 2011;39:1517
27. [56] Pai S, Ledoux WR. The shear mechanical properties of diabetic and non-diabetic plantar soft tissue. J Biomech 2012;45(2):364
70.

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[57] Hsu CC, Tsai WC, Hsiao TY, Tseng FY, Shau YW, Wang CL, et al. Diabetic effects on microchambers and macrochambers tissue properties in human heel pad. Clin Biomech 2009;24:682
6. [58] Gefen A. Plantar soft tissue loading under the medial metatarsals in the standing diabetic foot. Med Eng Phy 2003;25:491
9. [59] Corazza F, O’ Connor JJ, Leardini A, Parenti Castelli V. Ligament fibre recruitment and forces for the anterior drawer test at the human ankle joint. J Biomech 2003;36:363
72. [60] Pearsall AW, Kovaleski JE, Heitman RJ, Gurchiek LR, Hollis JM. The relationships between instrumented measurements of ankle and knee ligamentous laxity and generalized joint laxity. J Sports Med Phys Fit 2006;46(1):104
10. [61] Forestiero A, Carniel EL, Fontanella CG, Natali AN. Numerical model for healthy and injury ankle ligaments. Australas Phys Eng Sci Med 2017;40:289
95. [62] Kovaleski JE, Gurchiek LR, Heitman RJ, Hollis JM, Pearsall AW. Instrumented Measurement of anteriorposterior and inversion-eversion laxity of the normal ankle joint complex. Foot Ankle Int 1999;20:808
14. [63] Kovaleski JE, Hollis MJ, Norrell PM, Vicory JR, Heitman RJ. Sex and competitive status in ankle inversion-eversion range of motion of college students. Percept Mot Skills 2004;99:1257
62. [64] Lapointe SJ, Siegler S, Hillstrom H, Nobilini RR, Mlodzienski A, Techner L. Changes in the flexibility characteristics of the ankle complex due to damage to the lateral collateral ligaments: an in vitro and in vivo study. J Orthop Res 1997;15:331
41. [65] Ringleb SI, Dhakal A, Anderson CD, Bawab S, Paranjape R. Effects of lateral ligament sectioning on the stability of the ankle and subtalar joint. J Orthop Res 2011;29:1459
64. [66] Hubbard TJ. Ligament laxity following inversion injury with and without chronic ankle instability. Foot Ankle Int 2008;29(3):305
11. [67] Fontanella CG, Favaretto E, Carniel EL, Natali AN, Constitutive formulation and numerical analysis of the biomechanical behaviour of forefoot plantar soft tissue. Proc Inst Mech Eng, Part H: J Eng Med 2014;228:942
91. [68] Forestiero A, Raumer A, Carniel EL, Natali AN. Numerical approach to evaluate the interaction phenomena between foot and insole by means of a numerical approach. Proc Inst Mech Eng H 2014;229:003
9.

Chapter 27

Clinical Examination of the Foot and Ankle Kalyani Rajopadhye Sigvard T. Hansen Foot and Ankle Institute, Physical Therapy and Hand Therapy Clinic, Harborview Medical Center, University of Washington, Seattle, WA, United States

Abstract This chapter outlines the components of a methodical examination of the foot and ankle to elucidate the reasoning behind the selected modules and their value in reaching an accurate diagnosis while examining the foot and ankle. Components of the physical examination include demographics, vital signs, patient history, assessment of pain, visual observation/inspection, lower extremity alignment, radiographic examination, range of motion/flexibility/joint mobility, muscle strength, sensory testing, circulation, foot and ankle specific testing, footwear examination, and functional assessment. Finally, a consideration of future biomechanical research is presented.

27.1

Introduction

The clinical foot examination can take from a few minutes, if cursory, to half an hour or more if the examiner is being thorough. Depending on the purpose (presurgical planning, followup postsurgical, research study, etc.), the level of detail and the number of raters may vary. The following details are meant to be nearly exhaustive, and roughly in the order they would be collected, but it is unlikely that any one clinical exam would include all of these aspects.

27.2

Demographics

Clinical examination begins with recording age, sex, race, and ethnicity of the patient [1]. Thereafter, body mass index (BMI) is calculated from the subject’s height and weight as this can have multiple implications toward foot and ankle health and function [2 4]. BMI contributes significantly towards clinical decision making in managing patients with foot and ankle dysfunction [2,5].

27.3

Vital signs

Vitals signs include body temperature, pulse rate, respiration rate, and blood pressure. They are imperative to any clinical assessment, and should be undertaken at the beginning of each examination.

27.4

Patient history

Garnering a thorough and relevant history is essential when performing a clinical exam of the foot and ankle. The interview should begin with a detailed assessment of the subject’s chief complaint. Meticulous note is made of the site of symptoms, onset, frequency, duration, exacerbating factors, and relieving factors. Associated constitutional symptoms may be noted such as night sweats, fever, or weight loss, which may be related to an infection or neoplasm [6]. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00013-5 © 2023 Elsevier Inc. All rights reserved.

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Functional loss related to the chief complaint may take the form of pain, difficulty in walking, navigating uneven surfaces such as inclines or stairs, or even tolerating footwear. Patients may endorse related restrictions in participating in functional activities of daily living, as well as recreational activities and these should be noted. Comprehensive medical and surgical history are also elicited at this point. Relevant medical history regarding metabolic conditions, rheumatological, psychological/psychiatric, or inflammatory conditions is gleaned. Allergies are assessed and underscored for possible future treatment contraindications. The impact on the musculoskeletal system of diabetes is extensive and profound, and hence essential to document [7]. Any central or peripheral nervous system disorders can also have a bearing on lower extremity function, as can preexisting orthopedic conditions. Previous sprains or surgeries to the lower extremity are noted in addition to an inclusive history of any traumatic injuries. Medications such as oral steroids, pain medications (narcotic or otherwise), DMARDS (disease modifying antirheumatoid drugs), anticoagulants, antibiotics, as well as recreational drugs are noted. Several classes of drugs may have an impact on musculotendinous tissue, e.g., use of steroids have multiple implications in patient management, and are hence recorded meticulously. Fluoroquinolones have been implicated in causing tendinopathy and are thus important to record [8]. Smoking is essential to note due to the significant impact it can have on tissue and bone healing [7,9]. Alcohol-related health conditions are important to assess due to any relevant complications that they may create at the foot and ankle [10,11]. In the case of a postsurgical examination, detailed operative history is obtained. It is important to consider any weightbearing precautions, which may preclude testing of gait or full weightbearing of the involved lower extremity [12,13]. Detailed operative notes, and imaging studies are invaluable to analyze to obtain a clear status of the foot and ankle. The social history of the subject is also assessed. Occupational description and demand should be considered. Vocations involving physical labor may require a different approach to management vs those that are sedentary in nature. Baseline or predysfunctional activity levels are determined, mainly walking tolerance, and activities related to health-maintenance, exercise, and recreation. Injury or dysfunction can have a significant impact on ability to perform activities of daily living independently, and may require assistance from caregivers [14]. The individual’s home layout is taken into consideration to assess their mobility and impediments in accessing household areas such as stairs and bathrooms. Note is made of the presence of or need for assistive equipment such as walkers, crutches, or canes [14]. Immobilization/stabilization devices can take the form of controlled ankle motion (CAM) walker boots, splints, braces, and a variety of orthoses [14].

27.5

Assessment of pain

Pain is the most common presenting complaint at the clinician’s office and warrants a detailed evaluation [15]. Use of a body chart for symptoms can prove very useful in assessing the distribution of pain, and identifying referral patterns to reach an accurate diagnosis [14]. Nature or type of pain is helpful in identifying acuity, for example sharp versus dull, achy, or throbbing. Characteristics such as tingling, numbness, or burning suggest neurological involvement. Exacerbating and relieving factors affecting pain can offer insight into the etiology of pain. A pain scale of 0 to 10 is often useful.

27.6

Visual observation/inspection

27.6.1 Skin Any injuries to the skin are noted, including incisions. A thorough assessment of any wounds is required at this time. Any abnormal protrusions, exostoses, or lesions are noted. Pigmentations and skin lesions may indicate systemic pathology, or they may be benign. Obvious abnormalities in color may be evident and in need of assessment. This will often include bruising around the foot and ankle region that may be indicative of recent trauma that is consistent with the location of bruising, or even proximally. Either should prompt a thorough exam. Certain ecchymoses patterns such as at the plantar aspect of the midfoot are concerning or even pathognomonic for serious ligamentous injuries and should warrant immediate further investigation [16,17]. Redness of skin is usually associated with inflammation and will require attention, and infective pathology should be ruled out. If redness is noted in conjunction with shiny appearance of skin, loss of hair, or coolness at the foot, these

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symptoms may be indicative of an autonomic nervous system issue which is often observed in relation to prolonged immobilization and nonweightbearing episodes usually subsequent to significant traumatic injuries.

27.6.2 Edema Edema can be an important indicator of foot and ankle dysfunction. It may be localized to an area of the foot and ankle, or even generalized throughout the lower extremity, and it may be unilateral or bilateral. Presence of bilateral edema may be indicative of systemic pathology such as cardiac or renal disease [6] vs local edema which is often indicative of inflammation from localized pathology. Edema often results in increased pain, neural irritation, loss of range of motion, impacts wound healing, and ability to wear or tolerate footwear. Swelling or edema is a collective term used for an abnormal accumulation of excess fluid either intracapsular from hemorrhage or synovial fluid, or extra-capsular from extravasation of serum, blood, and lymph. While it is difficult to distinguish between the two, there may be some clear indicators as to the type of edema such as its localization. Several quantification methods have been used for edema. Determining the nature of edema is important: acute versus chronic, pitting versus non-pitting, localized versus generalized, location, etc. Presence of varicosities suggests compromised venous return. Long-term edema can create fibrinogen which might eventually convert to fibrin or scar tissue and create a pitting characteristic when pressure is applied to the skin. Such characteristics may offer valuable insight into managing edema. The figure-of-eight technique is one of the few methods described for the purpose of quantifying edema that has shown to be valid [18].

27.6.3 Atrophy Muscle atrophy (Fig. 27.1) may be evident in long standing disuse or nonuse of the extremity and may often extend to proximal muscles such as those of the thigh, and hips.

27.6.4 Temperature Warmth is usually indicative of inflammation, and coolness can occur due to autonomic dysfunction.

FIGURE 27.1 Right calf shows distinct muscle atrophy compared to left.

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27.6.5 Scarring Scars, surgical or otherwise, are assessed for their appearance, any abnormal thickness, or mobility (Fig. 27.2). Adherent scars can impact underlying soft tissue mobility, cause entrapment of nerves or tendons, and reduce joint range of motion.

27.6.6 Callus patterns Owing to its unique contact surface and weight bearing responsibilities, the plantar surface of the foot often reveals unique abnormalities (Fig. 27.3). Calluses are noted with regard to their location, size, and impact on patient’s function. They provide valuable insight into the abnormal pressure areas of the foot. Calluses over dorsal interphalangeal joints often occur in claw toes. Cavus feet often develop calluses under the first metatarsal and/or fifth metatarsal heads, as well as fifth metatarsal base. Callus under the second metatarsal head is indicative of transferred push off forces from

FIGURE 27.2 Scarring over the first metatarsals (left) and Achilles tendon (right).

FIGURE 27.3 Distinct regions of callus over the second, fourth, and fifth distal metatarsal heads.

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the first metatarsal onto the second [19]. Plantar surfaces of feet are uniquely vulnerable to breakdown, especially in the setting of neuropathy.

27.6.7 Exostoses Abnormal bony prominences are not uncommon and are noted as well as correlated with radiographs.

TABLE 27.1 Common foot deformities in hindfoot, midfoot, and forefoot, along with the primary plane of deformity. Deformity-hindfoot

Plane

Equinus

Sagittal

Varus

Coronal

Valgus

Coronal

Deformity-midfoot/forefoot

Plane

Varus

Coronal

Valgus

Coronal

Abduction

Transverse

Adduction

Transverse

Deformity-forefoot

Plane

Hallux valgus/bunion

Transverse

Hallux varus

Transverse

Hammer/claw toes

Sagittal

Bunionette

Transverse

Cock-up toes

Sagittal

FIGURE 27.4 Foot with claw and crossover toes.

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27.6.8 Ankle and foot deformities Obvious deformities at the ankle and foot are noted and characterized (Table 27.1). Especially important is equinus deformity which refers to a plantarflexed position of the ankle joint with difficulty dorsiflexing to neutral when the knee is in an extended position [19,20]. While it is most obviously noted in neuropathological conditions such as cerebral palsy, stroke, or spinal cord injuries, it very commonly presents in a more subtle form in neurologically normal individuals, and is referred to as gastrocnemius equinus. Gastrocnemius equinus is regarded as one of the most significant mechanical stressors to the foot and ankle and is thus an essential clinical assessment performed during all examinations of the foot and ankle [20,21]. Owing to its subtle nature, gastrocnemius equinus can often be overlooked during a clinical examination and thus requires a lower index of suspicion on the part of the examiner [20]. Toe deformities are also readily evident and should be described in detail (Fig. 27.4). Typical examples of these include bunions, hallux valgus, hammer toes, claw toes, or crossover toes.

27.7

Lower extremity alignment

Lower extremity alignment as a whole directly impacts that of the foot and ankle, and as such, is an integral part of assessment [22,23]. This regional interdependence created by one part of the lower quarter on another necessitates assessment of deformities in the musculoskeletal chain proximal to the foot and ankle [24]. Alignment abnormalities are evaluated in the coronal plane in the form of varus or valgus, the sagittal plane in the form of flexion or extension/recurvatum, as well as the transverse or axial plane in the form of torsion/rotation (Table 27.2). Assessment is made as to whether a deformity primarily exists at a joint, or morphologically in a bone, or in combination. Several diagnostic tools in addition to observation may be used to assess these abnormalities including weightbearing radiographs, magnetic resonance imaging (MRI) studies, or computed tomography (CT) scans. These can assist not just in characterizing the misalignment, but also in quantifying the degree of abnormality. Torsional deformities can be especially challenging to detect and quantify by physical examination alone [25,26]. It is essential also to determine the locus of torsion along the lower extremity bones in order that a correct treatment plan can be established. Lower extremity alignment has substantial implications on clinical decision making, especially surgical [27,28].

27.8

Foot posture or foot shape

The posture of the foot is analyzed both in nonweightbearing resting position as well as in weightbearing stance. Evaluating foot posture, especially weightbearing, is an integral part of assessment as it will guide management. Foot posture abnormalities are associated with pain and dysfunction, and impact quality of life [29,30]. Foot posture is classified as neutral, planus, or cavus based on the shape and height of the medial longitudinal arch of the foot, as well as the frontal plane alignment of the hindfoot and the transverse plane alignment of the forefoot.

TABLE 27.2 Common lower extremity deformities along with the primary plane of deformity. Deformity

Plane

Internal femoral torsion

Transverse

External femoral torsion

Transverse

Internal tibial torsion

Transverse

External tibial torsion

Transverse

Genu varum

Coronal

Genu valgum

Coronal

Genu recurvatum

Sagittal

Tibia varum

Coronal

Tibia recurvatum

Sagittal

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Nonneutral foot postures are associated with foot injuries. The incidence of each morphological type of foot posture has been described as a “bell-shaped curve.” In a review of 2047 diabetic feet, Ledoux et al. found 57% neutral, 24% cavus, and 19% planus feet [31]. Traditionally, anthropometric values such as arch index, arch height, and navicular drop have been used toward classification of foot type [32]. While there is not a standardized method to assess foot posture, the Foot Posture Index has been established as a reliable and validated tool to assess foot posture clinically [33,34]. Key radiographic measurements are considered valid and reliable for assessing foot posture and are commonly used [35,36].

27.8.1 Planus foot type Feet are considered as flat (planus) when there is a lowered medial longitudinal arch, and the alignment of the hindfoot attains a valgus position, with the forefoot in compensatory varus and abducted (Fig. 27.5). An obvious clinical sign is for too many toes to be visible on the lateral aspect when observing the foot in weightbearing posteriorly [37].

27.8.2 Cavus foot type A foot is described as a cavus when the medial longitudinal arch is abnormally high, and the alignment of the hindfoot attains a varus position, with the forefoot adducted (Fig. 27.6). A cavus type foot is observed frequently in neuromuscular disorders such as Charcot-Marie-Tooth (CMT) disease, poliomyelitis, or congenital disorders such as adult residual clubfoot [38]. Evaluating a cavus foot should focus on assessing each individual contributor to the overall deformity. This may include hindfoot varus with a more horizontal morphology of the subtalar joint, a narrow talocalcaneal angle, a hyperplantarflexed first ray, or metatarsus adductus, which overall will result in a more lateral loading foot when in weight bearing. Forefoot-driven hindfoot varus is often the result of a hyperplantarflexed first ray [38,39]. The subtle cavus foot which is present in neurologically normal adults is a milder version of the cavovarus foot and is often missed during examination but may contribute to dysfunction [40]. The peek-a-boo heel sign was described as an indicator of hindfoot varus which may even be present in a subtle cavus foot [41]. It is the appearance of the heel pad on the medial side of the foot when the foot is observed from the front. Forefoot position is noted, if in varus, valgus, or neutral. Claw toes are a common feature in a cavus foot deformity [38].

FIGURE 27.5 Bilateral planus foot type. Note valgus hindfoot position.

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FIGURE 27.6 Bilateral cavus foot type.

27.9

Limb length

Limb length asymmetry is an important concern whilst assessing the foot and ankle. There is significant association between the presence of limb length inequality and the presence of gait deviation [42]. Frequently, this may result in a biomechanical impact on the foot and ankle on either limb of the individual. There is agreement in the literature that a limb length discrepancy of greater than 2 cm is associated with musculoskeletal and gait pathology, and that greater than 5 cm can lead to long-term issues [43]. Limb length inequality can be anatomical, resulting from a structural bone deficit, or functional, wherein there are equal bone segments on either side but there is muscle lengthening or shortening, a mechanical axis alteration, a faulty gait pattern, or a joint contracture [42]. Foot and ankle deformity can be the cause of limb length discrepancy such as in the case of equinus deformity of the ankle. This may inform clinical decision making, particularly as it relates to surgical decisions in the foot and ankle [44]. It is not only essential to quantify the length disparity between the two limbs, but also to evaluate the site/deformity creating the imbalance. Limb length discrepancy may be assessed clinically as well as radiographically [42].

27.10 Radiographic examination A basic series of ankle and foot radiographs is obtained which includes: anterior/posterior (AP), mortise, and lateral ankle views and AP, lateral, and oblique foot views. Additional and specific views are taken based on the structure being evaluated and suspected pathology [45].

27.11 Range of motion/flexibility/joint mobility Range of motion testing is a key component in the biomechanical examination of the foot and ankle which enables the examiner to assess dysfunction, as well as clinical progress (Table 27.3) [46]. Limitation in range of motion can create movement dysfunction and impact gait and other functional activities. There is limited standardization of ROM measurements of the foot and ankle in the literature [47]. Owing to the presence of multiple bony segments at the foot and ankle complex, testing range of motion of individual joints is challenging and is usually a conjoint measurement contributed by several joints is used [47]. Sealey et al. demonstrated that in the face of ankle arthrodesis, compensatory hypermobility occurs at the subtalar and midfoot joints which may lead to arthroses [48]. Owing to such a lack of

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TABLE 27.3 Common lower extremity deformities, along with the primary plane of deformity. Standard range of motion measures in a clinical exam Ankle dorsiflexion with knee extended Ankle dorsiflexion with knee flexed Ankle plantarflexion Hindfoot eversion Hindfoot inversion First metatarsophalangeal joint dorsiflexion First metatarsophalangeal joint plantarflexion First interphalangeal flexion Lesser toe flexion-extension at metatarsophalangeal joints Lesser toe flexion-extension at interphalangeal joints

FIGURE 27.7 Maximum plantarflexion (left) and dorsiflexion (right) on radiographs of the ankle.

FIGURE 27.8 Physical measurement of dorsiflexion with knee flexed (left) and knee extended (right).

specificity of joint testing, it is important that interpretation of range of motion measurements be placed in a clinical context during an exam. While several techniques may be employed to evaluate ROM, visual appraisal, and goniometric measurements are the most commonly used methods of assessment (Fig. 27.7). Recently, app-based goniometers have been increasingly used to assess these measurements [49]. Clinicians also routinely utilize radiographic measurements for assessment when these are available (Fig. 27.8).

27.12 Joint mobility Arthrokinematics of the joints is assessed within each individual joint based on its shape and capsular patterns [46]. In addition to the foot and ankle joints, superior and inferior tibiofemoral joint mobility is also tested due to its close proximity and functional alliance with the foot and ankle. Of particular importance in examination of the foot and ankle is testing addressed toward hypermobility of the first ray [19]. This has been implicated in a wide range of common foot

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TABLE 27.4 Injured ligament and corresponding test. Ligament

Test

Anterior talofibular ligament

Anterior drawer test

Calcaneofibular ligament

Talar tilt test

Syndesmotic ligament, anterior tibiofibular ligament, posterior tibiofibular ligament

External rotation stress test

Interosseous talocalcaneal ligament

Anterior drawer test

pathologies such as hallux valgus, metatarsalgia, metatarsal stress fractures and joint disruption, hammer toes, acquired flatfoot deformity, and tibialis posterior tendon dysfunction [50]. First ray mobility is tested clinically, and is typically also evident in weight bearing radiographs of the foot. It may also be assessed using a mechanical device such as that described by Klaue et al. [51].

27.13 Ligamentous/stability testing Ligamentous injuries to the ankle are one of the most common musculoskeletal injuries [52]. Lateral, medial, and rotational stability of the ankle most often require assessment due to disruption of ligamentous, or tendinous soft tissues with or without bony injuries (Table 27.4). There is an association between acute ankle sprains and future incidence of sprains. Additionally, disability can be related to long-term adverse outcomes such as chronic ankle instability and posttraumatic osteoarthrosis which is not uncommon as a sequela of instability related to ligamentous compromise [52]. Any history of twisting or sprain, recent or remote, should therefore prompt a thorough testing of ligamentous structures around the foot and ankle. Energy/mechanism of the injury, position of the foot, presence of rotational forces on the foot and ankle will guide the directional stress testing of tissues accordingly [53]. The inversion mechanism of injury is the most common mechanism encountered in the clinic and emergency room with subsequent injury to the lateral ankle ligaments, most often the anterior talofibular ligament and thereafter the calcaneofibular ligament. Associated peroneal retinaculum and peroneal tendon injuries should also be ruled out [54]. There is a significant association with the cavus type foot and lateral ankle injuries related to inversion [55]. There are other foot ligamentous injuries besides inversion sprains. Subtalar ligament sprains are not uncommon; however, they are commonly missed and may lead to ongoing dysfunction [56]. The interosseous talocalcaneal ligament is most commonly affected, after the calcaneofibular ligament [56]. Additionally, external rotation forces while the foot is in varus can create disruption of the deltoid, as well as syndesmotic ligaments and requires testing. These injuries are more commonly associated with fractures [57]. Furthermore, stress testing of the syndesmosis is necessary with high ankle sprains, with a pain distribution that extends from the anterior ankle proximally into the distal leg.

27.14 Tendon Tendon disorders are routinely encountered in the foot and ankle clinic. A common injury is the rupture of the Achilles tendon, which leads to the immediate inability to plantarflex the foot during gait and other activities. A positive Thompson’s test, palpable defect at the area of the Achilles tendon, and inability to produce plantarflexion force prompt immediate medical attention, and further diagnostic imaging for confirmation of the rupture. Traumatic and atraumatic ruptures of the tibialis anterior tendon may also occur, the former usually through a forced plantarflexion mechanism, lacerations and the latter in the setting of degenerative tendinopathy through chronic overuse and gradual weakening of the tendon [58]. Generalized joint hypermobility can have a significant impact on joint stability, and despite having a high prevalence in the population, is often overlooked [59]. Frank hypermobility syndromes such as Ehler’s Danlos syndrome not only contribute significantly to instability but are also associated with chronic musculoskeletal pain and fatigue [59]. Degree of hypermobility can be assessed and scored using the Beighton Index [59].

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27.15 Muscle strength Muscle strength is often tested grossly, but specific muscle testing is also important. Hand-held dynamometry is a reliable and validated method to assess muscle strength [15]. Manual muscle testing is the most widely performed technique to assess muscle strength, especially in clinical settings where there may not be access to equipment. The disadvantages of this method is that it can lack sensitivity at higher grades, and can be dependent on the strength of the examiner. Group muscle testing as well as specific muscle testing are required while clinically examining a patient. Group muscle testing may be required for the entire lower extremity to test for deconditioning in the face of prolonged lower extremity nonuse, or to identify any proximal muscle weakness impeding the gait cycle. This includes muscle testing of musculature controlling the knee, the hip, and even that involved in lumbopelvic control. Specific muscle testing is performed to assess the following muscles (Table 27.5). Specific muscle testing may be selected based on suspicion related to a diagnosis such as myotomal testing in the case of lumbar radiculopathy, or neural injury. Intrinsic muscle testing is also essential to assess, and becomes especially important when assessing midfoot, and forefoot issues [60]. Group testing of intrinsic muscles has been shown to be reliable and valid [61].

27.16 Sensory testing A 10 g monofilament is recommended for sensory testing as it provides a threshold for sensory protection of the soft tissues of the foot [62], and is a sensitive as well as specific test as regards to superficial and deep pressure sensation (Fig. 27.9). Lack of sensation to the Semmes-Weinstein 5.07 filament is a risk factor for ulceration in diabetic feet [63]. Vibratory sensation may also be tested with a tuning fork [62]. Jeng et al. studied sensation in 40 young adults indicating that an inability to feel a 5.07 monofilament would indicate a loss roughly of 98% of sensory capability [64]. Traumatic injuries of the foot and ankle often result in peripheral neuropathies and necessitate testing of individual nerve distributions (Table 27.6). Neural traction injuries are common and may present with hyperesthesia in the respective nerve distribution [65].

TABLE 27.5 Specific muscles that are often strength tested. Key muscles of the foot and ankle Tibialis anterior Tibialis posterior Peroneus longus Peroneus brevis Extensor hallucis longus Extensor digitorum longus Flexor hallucis longus Flexor digitorum longus Gastrocnemius and soleus

FIGURE 27.9 A 10 g monofilament (left) and sensory test being performed.

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TABLE 27.6 Key nerves of the foot and ankle. Superficial peroneal Deep peroneal Sural Saphenous Tibial Medial plantar Lateral plantar

TABLE 27.7 Special tests for the foot and ankle [6,67,68]. Test

Indication/tissue being tested

Silfverskiold test

Gastrocnemius equinus

Thompson test

Achilles tendon rupture

Flexible versus rigid flat foot

Flat foot characteristics

Babinsky’s sign

Upper motor neuron disease

Oppenheim’s test

Upper motor neuron disease

Tinel’s test

Neuropathy

Dorsiflexion-eversion test

Tarsal tunnel syndrome

First ray hypermobility

First ray instability

Mulder test

Neuroma

Coleman block test

Forefoot-driven hindfoot varus

Well’s criteria

Deep vein thrombosis

27.17 Circulation Capillary refill testing is performed along with tibial artery and dorsalis pedis pulses [6]. Femoral and popliteal pulses may need to be examined in addition to local foot pulse monitoring [66].

27.18 Foot and ankle specific testing Diagnostic accuracy of structure specific tests in the foot and ankle is not yet fully established in the literature. However, several tests (Table 27.7) are performed routinely in the clinic based on diagnostic suspicion [67].

27.19 Footwear examination Footwear has a significant impact on the biomechanics of the foot and ankle [69]. Incorrect size and style of footwear can contribute to or even create pain, deformity, and dysfunction [69]. As such, a thorough footwear exam is an essential component in assessing the biomechanics of the foot and ankle. The components of a footwear exam should include assessment of wear patterns, cushion, midsole structure, overall support level, heel-toe drop, forefoot width and volume, and heel flare. This will reveal overall compatibility of the footwear, or the lack thereof with the patient’s foot. Presence of neuropathic disorders such as diabetes necessitates an especially thorough and specific evaluation of footwear.

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TABLE 27.8 Functional tests in a foot and ankle examination. Gait Single leg balance testing Squatting Stair step ascent, descent Incline decline assessment

27.20 Functional assessment This allows for assessment of compound movements in functional activities to examine intersegmental movement and control. Functional movements are able to engage several faculties in conjunction, such as range of motion, strength, coordination, and balance, and can reveal abnormalities that may be missed on other exams [66,70]. Gait is perhaps the most central of functional assessments as regards examination of the foot and ankle (Table 27.8).

27.21 Outcomes assessment In the past few decades patient-reported outcomes assessments have become an integral part of evaluating healthcare giving the patient their rightful voice in their healthcare management. Outcomes assessments guide patient choice, boost communication and interaction between the clinical provider and patient, inform economic policy, healthcare policy, study populations, improve quality of health services, and provide evidence of health and need in groups. They are an essential component of research that assesses effectiveness [71 73]. Several instruments/questionnaires have been developed for various generic health conditions, as well as more regionally specific ones for the foot and ankle [74]. Within the foot and ankle, disease-specific outcome measures are frequently used such as the Ankle Osteoarthritis Scale [74]. Reliability, validity, responsiveness, and feasibility are important psychometric properties in outcome measures that need to be established prior to their use in clinical or research settings [71]. Most foot and ankle publications use a generic, as well as a foot and ankle specific outcome measure [74]. Generic outcome measures commonly used in the foot and ankle literature and clinical practice are the Visual Analog Scale, the Numerical Pain Rating Scale (NPRS), and the SF-36 scale [74]. The Visual Analog Scale is one of the most commonly used outcome measures in the foot and ankle literature for the assessment of pain. While it has demonstrated good sensitivity and accuracy for acute pain, these properties have not been established for chronic pain in the outpatient setting [75]. The NPRS is also a widely used and validated tool in the assessment of pain [76]. In conjunction with the Visual Analog Scale, the SF-36 survey is the other most commonly used nondisease-specific outcome measure utilized in foot and ankle publications, and has demonstrated adequate clinimetric properties [77]. Despite being the most widely used in the foot and ankle literature, use of AOFAS outcome score is not supported as a sole outcome tool due to inadequate reliability and validity [73,77,78]. The Foot and Ankle Ability Measure (FAAM) was created to detect any changes in physical function in patients with leg, foot, and ankle musculoskeletal disorders and includes an activity of daily living subscale and a sports subscale [73]. While the FAAM has good psychometric properties for assessing individuals with a broad range of musculoskeletal disorders of the lower leg, foot, and ankle [73,79], its efficacy needs to be established for patients with specific foot and ankle conditions [80]. The Manchester-Oxford foot questionnaire has been shown to have promising psychometric properties toward assessment of this population [81]. The Lower Extremity Functional Scale has also been validated, although it does display a high ceiling effect [82]. The Ankle Osteoarthritis Scale was the most commonly used disease-specific outcome measure in the foot and ankle literature [74]. The Foot and Ankle Disability Index has been shown to have good clinimetric properties for patients with chronic ankle instability [77,80]. The advent of computer adaptive testing has allowed for improved ease of administration, and accuracy of testing for all conditions. The Patient-Reported Outcomes Measurement Information System (PROMIS) is one such test that has been validated for a variety of foot and ankle conditions [73]. It is a computer adaptive test with several domains that has been shown to be reliable, valid, and responsive for several foot and ankle conditions [73,79,83]. PROMIS has shown to have better accuracy and a stronger correlation with preoperative and postoperative function than the NPRS [76].

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Use of patient-reported outcome measures has its limitations, such as their inability to differentiate why there is a limitation in a task, inability to quantify the impact that a limitation has, nor a change in impact. These tools are also dependent on the subject’s sound cognition, judgment, and willingness and ability to complete the outcome measure. There may be limited correlation between the physical performance measures and scores. Owing to limitations of each outcome measure, it is not recommended to use a single outcome tool in the research setting for the foot and ankle patient population [73].

27.22 Areas of future biomechanical research There are several prominent areas in need of further biomechanical research at this time with wide-ranging implications toward our understanding of the foot and ankle and clinical care. Technological advances such as motion capture technology, pedobarographic measurement technology, machine learning, etc. have allowed for broadening the scope and precision of biomechanical research in the vast and complex research area that is the foot and ankle [15,84]. There are a few prominent areas of research to highlight within this field. Studying gait offers a way to assess and improve our understanding of the kinetics and kinematics of the lower extremity and offers many opportunities in understanding dynamic vs static relationships. While gastrocnemius equinus is widely recognized as a source of pathology at the foot and ankle, no validated measurement techniques have been identified. Given the overarching impact this entity has on the mechanics of the foot and ankle, inability to reliably measure it creates a barrier to recognizing and managing equinus. Additionally, while the testing of gastrocnemius equinus is currently performed in nonweight bearing positions, there is a need to understand the biomechanical effect of equinus forces at the foot and ankle during functional gait. These limitations are also applicable to analyzing ideas such as first ray hypermobility. Another area of research is establishing a reliable and valid method to measure lower extremity torsion. This has become of considerable importance, especially with the advent of total ankle replacement techniques, and the ensuing need to surgically position the prosthetic implant in an optimal alignment. Given that foot posture is a construct that is considered clinically for nearly every patient being assessed for the foot and ankle, and although there are some validated measurements to classify foot posture, this area still needs a fair amount of attention with regards to biomechanical research. Evolving techniques in various clinical fields related to management of foot and ankle dysfunction such as surgery, physical therapy, prosthetics, and orthotics all require the support of biomechanical research in the future. Overall, the level of evidence continues to be low in the literature with suboptimal methodology, although this increases with time [85]. While the majority of the publications in the foot and ankle literature are therapeutic in nature [85,86], there is certainly room for more quality research in predictive modeling, diagnostics, and social aspects of patient care.

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[48] Sealey RJ, Myerson MS, Molloy A, Gamba C, Jeng C, Kalesan B. Sagittal plane motion of the hindfoot following ankle arthrodesis: a prospective analysis. Foot Ankle Int 2009;30(3):187 96 Epub 2009 Mar 27. [49] Wang K, Hussaini S, Teasdall R, et al. Smartphone applications for assessing ankle range of motion in clinical practice. Foot Ankle Orthop 2019;4(3). Available from: https://doi.org/10.1177/2473011419874779. [50] Glasoe WM, Michaud TC. Measurement of dorsal first ray mobility: a topical historical review and commentary. Foot Ankle Int. 2019;40 (5):603 10. Available from: https://doi.org/10.1177/1071100719839692. [51] Shibuya N, Roukis T, Jupiter D. Mobility of the first ray in patients with or without hallux valgus deformity: systematic review and meta-analysis. J Foot Ankle Surg. 2017;56(5):1070 5. [52] Herzog M, Kerr Z, Marshall S, Wikstrom E. Epidemiology of ankle sprains and chronic ankle instability. J Athl Train. 2019;54(6):603 10. 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Chapter 28

Foot Type Biomechanics Scott Telfer1,2,3 and William R. Ledoux1,2,3 1

Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States, 2RR&D Center for Limb Loss and MoBility

(CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 3Department of Mechanical Engineering, University of Washington, Seattle, WA, United States

Abstract The morphology and function of the human foot vary significantly across individuals, generally a result of the underlying musculoskeletal structures. While these are continuous, multi-dimensional spectrums, for clinical and anatomical ease of categorization, feet are generally grouped into three types: cavus (high arched), neutral, or planus (flatfoot). This chapter provides an overview of these foot types, reviewing their structural and functional differences, as well as the tools that are used to help define each type. An overview of the biomechanics of different foot types will be presented. Associations between foot type and clinical problems along with potential treatments will be covered. Future areas of research will also be discussed.

28.1

Introduction

Feet come in many different shapes and sizes, and these differences may affect the way they function. Moreover, the shape and function of the foot can change over the course of an individual’s life. Factors including genetics [1], footwear [2], environment, and disease may all play a role in the development of and changes to the structure and function of the foot. Clinically, feet are generally categorized into three types (Fig. 28.1): cavus (high arched), neutral, or planus (flatfoot). Loosely defined, cavus feet generally have an inverted hindfoot (coronal plane), a high arch (sagittal plane), and an adducted forefoot (transverse plane), while planus feet have an everted hindfoot, low arch, and an abducted forefoot; neutral feet are well aligned in all three planes. These definitions are really continuums and it is difficult to classify what is “normal” or “healthy,” in particular when pain is considered. For instance, a completely flat foot could also be pain free, while a neutral, well-aligned foot might be painful. Weightbearing is another important aspect—a foot might have a normal arch when unloaded, but it could collapse when weight bearing. Functionally, these foot types are closely related to what are often referred to as “supinator,” normal, or “pronator.” As discussed elsewhere in this book, the terms supination and pronation are often poorly defined in relation to their specific biomechanical characteristics in the lower extremity, so we are avoiding their use. Cavus and planus foot types have been suggested to be associated with a range of musculoskeletal problems, both within the foot and more proximally (discussed later). Several tools have been presented to quantify foot type, and many of which are discussed later in this chapter. However, there is no true universally accepted gold standard for categorizing different feet, and that foot type itself is a clinical concept used to simplify the variation and complexities of the overall anatomy. As noted, cavus feet typically feature a raised arch, plantarflexed first ray, forefoot adduction, and an inverted hindfoot [3]. Dynamically, we would expect these feet to display hindfoot inversion compared to neutral feet during gait. This deformity has a prevalence of around 10% 15% in the general population [4]. People with pes cavus are reported to have a higher prevalence of foot pain than those with neutral feet [5]. While factors such as muscle weakness caused by peripheral neuropathy (often secondary to diabetes or Charcot Marie Tooth disease), trauma, or the congenital effects of club foot are thought to play a role in the development of the cavus foot, a considerable proportion of cases are considered idiopathic [6]. Further, planus feet typically present with a lowered arch and everted hindfoot when compared to the neutral foot, and dynamically would be expected to display excess hindfoot eversion during walking or running. This is a relatively widespread deformity, with a prevalence of around 20% in the general population [7], although this value may vary depending on how the foot type is defined. Flexible planus feet are commonly seen in infants and young children, but the majority will develop a normal medial arch by age 5 or 6. In adulthood, factors that have been associated with Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00043-3 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 28.1 Anatomical models showing the structure and shape of the foot for the three main foot type classifications. Note in particular the relative changes in bone position of the medial longitudinal arch. Top: cavus; middle: neutral; bottom: planus.

flatfoot include age, female sex, and body mass index [8]. One of the most common causes of acquired planus foot type during adulthood is posterior tibial tendon dysfunction, which affects one of the main structures that support the medial side of the foot [9], initiating the collapse of the arch. For individuals who do have clinical symptoms associated with cavus or planus foot types, there are a range of treatment modalities available. More conservative options include foot orthoses to provide external support to the foot, or strengthening and stretching exercises to develop the intrinsic muscles. For more severe cases that lead to pain and/or functional limitations, surgical corrections can be performed.

28.2

Structural foot type

A variety of static measures have been proposed to aid in quantifying foot structure, some based on clinical measurements of morphology, others on medical imaging, and some on the plantar footprint. Some of the most commonly used measures are described in the following sections. For additional measures, we refer readers to relevant review articles in the literature (for example [10,11]). X-ray imaging: Weight-bearing radiographic views of the foot and ankle can be used to assess foot type. Sagittal plane measurements including (but not limited to) the lateral talometatarsal angle (Meary’s angle), talocalcaneal angle, and calcaneal pitch angle are often used (Fig. 28.2). In many cases, the flexibility of the foot means that the amount of loading on it can affect these measurements, however it has been shown that loading greater than 20% [12] or 25% [13] bodyweight is in most cases sufficient to stabilize these values. Most measures of foot type are made in the sagittal plane and are related to arch height. However, in the transverse plane, one common X-ray measurement is the talonavicular coverage angle [14], which accounts for the rotation of the forefoot relative to the hindfoot. In the frontal plane, the hindfoot alignment view can be used to quantify the position of the calcaneus relative to the tibia [15]. CT imaging: Previously, partial weightbearing CT scans have been used to quantify foot morphology, although bone coordinate systems, particularly in the hindfoot, defined by internal matrices were not very sensitive to foot type [16]. In recent years, weightbearing CT scanners designed specifically for imaging the foot and ankle have become more accessible and clinically available. It has been suggested that, through the use of similar measurements to those used in assessing foot type for standard X-rays, as well as novel measurements taking advantage of the 3D nature of the data, these weightbearing scans can more clearly show the changes in bony alignment in patients with planus feet [17], and may provide a more accurate assessment of flexible planus deformity than standard X-rays [18].

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FIGURE 28.2 Examples of some commonly used radiographic measurements on a lateral X-ray of a planus foot. From Look N, Autruong P, Pan Z, Chang FM, Carollo JJ. Radiographic and plantar pressure assessment of pes planovalgus severity in children with cerebral palsy. Clin Biomech 2021;85:105364.

FIGURE 28.3 The arch index. From Cavanagh PR, Rodgers MM. The arch index: a useful measure from footprints. J Biomech. 1987;20:547 51.

Foot posture index (FPI): The FPI is a six-item reference tool based on visual observations of the hindfoot and forefoot with the subject in relaxed standing [19]. Items are scored on a five-point scale (22 to 2) and summed, with a significantly negative score indicating a cavus foot, and positive indicating planus. The tool is generally considered to be valid and reliable, although some questions of its reliability for use in older adults have been raised [20]. Navicular height and navicular drop: Navicular height is measured with the patient in relaxed standing. The most medial prominence of the navicular is palpated and marked. A vertically placed ruler is then used to measure the height of the tuberosity from the ground. Modifications including normalizing to foot length have been proposed [21]. Navicular drop is a variation on this measurement where, again with the patient standing, the foot is adjusted by the rater to subtalar neutral position, the navicular height measured, then the foot allowed to drop to its relaxed position, and the navicular height measured again. The difference between the heights is the output measurement, with planus foot type expected to show larger values. This measurement has been found to have only moderate reliability [22]. The arch index: A footprint is obtained, either using a pressure measurement platform or inked paper, and the measurement is calculated as the ratio between the contact area of the middle third of the footprint and the overall area of the footprint (Fig. 28.3) [23]. The accuracy of footprint-derived estimates of foot type is influenced by body composition and may have limited reliability [24].

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Arch height index (AHI): The AHI is calculated by dividing the dorsal height of the foot, measured at half of the total foot length, by the truncated foot length, defined as the length of the foot from the posterior of the heel to the first metatarsal head. A measurement system is available to assist in obtaining these values [25].

28.3

Functional foot type

Quantifying the dynamic behavior of the foot can also be used to define different foot types. Indeed, several studies have reported that static measurements do not necessarily correlate well with the functional behavior of the foot [26,27]. It is likely that dynamic variations in foot function during activities where loads are greater than in relaxed standing and neuromuscular control is challenged are likely to be more closely associated with clinically relevant problems. Dynamic arch index: Dynamic measurements of the arch index, calculated in the same manner as described earlier for the static arch index except using a footprint obtained during walking rather than relaxed standing, are more reliable than static arch index measurements [28]. Center of pressure excursion index (CPEI): The CPEI is also calculated using a dynamic pressure footprint. It is defined as the concavity of the center of pressure line at the metatarsal head region, normalized to the foot width at that region [29]. Planus feet tend to show a smaller CPEI, cavus feet a larger value [30]. Dynamic navicular drop: Navicular drop (described earlier) can be measured during dynamic activities such as walking or running gait with the aid of a motion capture system and a marker placed on the navicular tuberosity (Fig. 28.4). Measurements of dynamic navicular drop are not closely associated with static measures [26], although skin motion artifact can cause the amount of drop to be underestimated [31].

28.4

Foot type biomechanics

Given the underlying variations in the morphology and structure of different foot types it is not surprising that their biomechanics also vary; indeed, as discussed in the previous section these changes can be used to define different foot types. This section discusses some of the notable differences found between foot types. The distribution of ground reaction forces (GRFs) differ between foot types, as planus feet have demonstrated more loading under the first metatarsal compared to neutral feet. Arch index has also been associated with positively with midfoot loading, that is, flatter feet have more midfoot loading [32]. Queen et al. explored differences in plantar loading between flat and neutral feet during athletic tasks, and found that flat feet have different loading patterns [33]. However, when studying healthy female runners, it was shown that the vertical GRF profiles are similar [34]. Plantar pressures have significant differences between foot types. Examining 400 healthy subjects, SanchezRodriguez measured FPI and plantar pressure [35]. They found decreased loading under the hallux and lesser toes, and increased loading under the fifth metatarsal head for cavus feet; the exact opposite relationship was seen for planus feet. In a study of 92 individuals, Buldt et al. [36] found that different foot postures, defined by the FPI and other static measures, had unique plantar pressure characteristics (Fig. 28.5). A review of the literature by the same group found that for planus feet there was evidence of (amongst other variables) elevated peak pressure and pressure time integral in the medial arch, central forefoot, and hallux. Cavus feet had elevated pressures at the heel and lateral forefoot [37].

FIGURE 28.4 The dynamic navicular drop height measurement obtained from (A) a marker on the navicular tuberosity, and (B) using an ultrasound probe attached to the foot. M1TH, Medial first metatarsal head; NAV, navicular; SUST, sustentaculum tali. From Telfer S, Woodburn J, Turner DE. An ultrasound based non-invasive method for the measurement of intrinsic foot kinematics during gait. J Biomech. 2014;47:1225 8.

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FIGURE 28.5 Top: example peak pressure diagrams for foot posture groups; bottom: comparisons of peak pressure for all foot posture groups. For the foot posture comparisons, red indicates larger peak plantar pressure values and blue indicates lower peak plantar pressure values. From Buldt AK, Forghany S, Landorf KB, Levinger P, Murley GS, Menz HB. Foot posture is associated with plantar pressure during gait: a comparison of normal, planus and cavus feet. Gait Posture 2018;62:235 40.

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FIGURE 28.6 Average hindfoot kinematics in major anatomical planes of motion for different foot type groups during walking. Rectus, neutrally aligned. From Kruger KM, Graf A, Flanagan A, McHenry BD, Altiok H, Smith PA, Harris GF, Krzak JJ. Segmental foot and ankle kinematic differences between rectus, planus, and cavus foot types. J Biomech 2019;94:180 186.

A systematic review of studies that assessed kinematic differences between foot types (Fig. 28.6) found that planus feet were associated with increased frontal plane motion of the hindfoot during gait [38]. There was little evidence of significant effects on more proximal joints. Another systematic review of the literature found moderate evidence for an association between foot posture and subtalar joint kinematics and leg stiffness during running, although the authors noted that the results were limited by the heterogeneity of study designs [39]. Using a five-segment foot model, one group explored how foot kinematics varied between cavus, neutral, and planus subjects [40]. Feet were group based on based on normative data for the FPI, AI, and normalized navicular height. Several kinematic differences were determined, including that cavus feet have altered frontal and transverse hindfoot motion, that cavus feet have less midfoot motion during the first half of stance, and that planus feet have less midfoot frontal plane motion. Individuals with planus or cavus feet have poorer postural control than neutral feet [41]. In this study, the authors tracked center of pressure during single limb stance with eyes closed in three groups with different foot types and found the cavus group to have greater deviations in all directions, and the planus group to have greater deviations in the anterior posterior direction. Other groups have shown that postural stability is decreased with cavus feet [42]. Muscle activity during gait for different foot types has also been studied [43]. Significant differences were found, particularly for peroneus longus and tibialis anterior. Foot type has also been found to influence the activity of more proximal muscles, with gluteus medius found to have increased activity in cavus feet compared to normal and planus feet [44]. Bone morphology varies due to foot type, changes that may be related to different loading patterns between foot types [45,46]. Metatarsals of cavus feet were found to have small cross sectional areas [45], while the calcanei of planus feet had decreased height and increased length and the tali of cavus feet demonstrated increased posterior mass [46]. Others have found differences in the talar and navicular articular surfaces for planus feet [47].

28.5

Association with pain and injury

28.5.1 Pain and injury Cavus and planus foot types have traditionally been considered to be related to pain and injury risk. This is perhaps best represented historically by the disqualification of those with planus feet from military service [48]. More recent literature is generally supportive of this relationship, although there is some conflicting data. In addition, much of the research in this area is of limited quality, and affected by heterogenous or highly subjective measures of foot type. There are a few large, epidemiological studies that give some insight into the association between foot type and clinical problems. The Cheshire Foot Pain and Disability Study found that both cavus and planus foot type were associated with self-reported disabling foot pain [49], although follow-up clinical assessments did not reproduce this finding. A study in Denmark found a similar association between self-reported cavus and planus foot type and foot pain [50]. One of our major sources in this area is the Framingham foot study, a large epidemiological study that included measures of foot function (plantar pressures), structure, and a range of clinical outcomes from .3000 subjects, which has produced several publications. This work has provided insights into the relationship between foot pain and posture. Using static pressure measurements to categorize foot type, this study reported that planus foot type was associated

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with arch pain (odds ratio 1.38). Cavus foot type was also found to have a protective effect against pain in the ball of the foot (odds ratio 0.74) and the arch (odds ratio 0.64) [51]. In another study based on the Framingham foot dataset, it was found that planus foot type was associated with increased likelihood of having hammer toes (1.38) and overlapping toes (1.44) [52]. Furthermore, those with planus foot posture had a 78% higher risk of multiple falls [53]. Results from the Framingham Foot Study found that feet which were “pronated” in function were associated with back pain (odds ratio 1.51) [54]. Military cohorts have also been a valuable source of information on how foot type may relate to clinical conditions, although the demographic make-up of these cohorts may limit the external validity of the findings. Planus foot posture has been found to be associated with foot and ankle injury risk in military cadets [55]. Those with cavus foot types are at greater risk of developing metatarsal stress fractures [56]. Trainees with planus feet are more likely to develop lower extremity stress fractures compared to neutrally aligned feet [57] and of the 35% of naval recruits that developed stress fractures over a 10-week basic training, more than 70% had planus feet [58]. In terms of disease, individuals with diabetes are more likely to have pes cavus foot type than the general population, and this foot type was associated with a high incidence of other foot deformities including hammer/claw toes and prominent metatarsal heads [59]. It was also demonstrated in a well-characterized, high-risk diabetic population of 398 subjects that planus feet were associated with hallux valgus and hallux rigidus, and the cavus feet were associated with hammer/claw toes [60]. Knee osteoarthritis is also thought to be associated with foot and ankle biomechanics that are seen in planus feet [61].

28.6

Treatments

Foot orthoses, which can vary widely in terms of form and material, have differing effects on planus feet compared to normal feet, in terms of kinematics and kinetics [62], as well as plantar pressures and muscle activity [63]. These are appealing as a treatment option given their low cost and ease of use. Exercise programs that target the intrinsic muscles of the foot have been proposed to help treat symptomatic planus feet. Only a few small randomized trials comparing intrinsic foot muscle exercise program to usual care have been carried out in this area, but these preliminary results have been positive, with the treatment groups seeing significant improvements in pain, navicular drop, and foot posture index [64,65]. Other treatments including taping have been investigated, finding short-term improvements in navicular height although these effects were found to reduce after exercise [66]. In cases with severe pain and/or deformity leading to functional limitations, surgical interventions may be required to realign the anatomy and improve the functional biomechanics of the foot. The indications and techniques for these are discussed in detail in dedicated chapters in this book, and the reader is referred there for more information.

28.7

Areas of future biomechanical research

While we have a relatively good understanding of foot posture, extensive gaps in our knowledge remain. Recommendations for which tools to use to define different foot types and under which circumstances, understanding mechanisms of injury relating to foot type, and understanding how different foot types develop would be beneficial. Novel technologies such as Dynamic 3D scanning (Fig. 28.7), where the external surface of the foot is captured and represented as a 3D model, may provide insights into foot function [67]. The ability to measure how the overall foot deforms during dynamic activities may provide a more detailed look at how the different foot types vary in function. There may be a role for patient-specific computational simulations of foot function to help understand the development of different foot types, understand the changes in intrinsic loading, and optimize surgical corrections. In children, it remains challenging to differentiate between pes planus as part of the natural course of development and those who have clinical problems that require treatment [68]. Improving our ability to identify those cases which will require intervention could allow for earlier and more effective treatment. Elucidating the mechanisms behind cavus feet currently defined as idiopathic would also be valuable. The role of intrinsic muscles in different foot types is also worth further investigation [69]. In terms of treatments, programs for strengthening the intrinsic muscles show some promise, and further investigation with larger randomized controlled trials to determine if these programs can maintain improvements in foot posture and function over longer periods of time and across different groups of patients are warranted.

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FIGURE 28.7 Results from 3D dynamic foot scanning. Inclusion of points lower than a height threshold of 15 mm for the next data processing (A) “red-yellow-blue” color scale mapped height values close to 0 mm in red, close to 15 mm in blue; (B) vertical projection of these points on the ground after morphological opening; (C) same processing on each frame of the stance phase; (D) LHDs picture summarizing the whole foot sole pictures of the stance. From Samson W, Van Hamme A, Sanchez S, Che`ze L, Jan SVS, Feipel V. Foot roll-over evaluation based on 3D dynamic foot scan. Gait Posture 2014;39:577 82.

References [1] Telfer S, Bigham JJ, Sudduth ASM. Plantar pressures in identical and non-identical twins. J Biomech 2019;86:247 50. ˆ t K, Pataky TC, De Clercq D, Aerts P. The effects of habitual footwear use: foot shape and function in native barefoot walkers. [2] D’AoU Footwear Sci 2009;1:81 94. [3] Younger ASE, Hansen ST. Adult cavovarus foot. J Am Acad Orthop Surg 2005;13:302 15. [4] Sachithanandam V, Joseph B. The influence of footwear on the prevalence of flat foot. A survey of 1846 skeletally mature persons. J Bone Jt Surg Br 1995;77:254 7. [5] Burns J, Crosbie J, Hunt A, Ouvrier R. The effect of pes cavus on foot pain and plantar pressure. Clin Biomech 2005;20:877 82. [6] Burns J, Crosbie J. Weight bearing ankle dorsiflexion range of motion in idiopathic pes cavus compared to normal and pes planus feet. Foot 2005;15:91 4. [7] Dunn JE, Link CL, Felson DT, Crincoli MG, Keysor JJ, McKinlay JB. Prevalence of foot and ankle conditions in a multiethnic community sample of older adults. Am J Epidemiol 2004;159:491 8. [8] Pita-Ferna´ndez S, Gonza´lez-Martı´n C, Seoane-Pillado T, Lo´pez-Calvin˜o B, Pe´rtega-Dı´az S, Gil-Guille´n V. Validity of footprint analysis to determine flatfoot using clinical diagnosis as the gold standard in a random sample aged 40 years and older. J Epidemiol 2015;25:148 54. [9] Yao K, Yang TX, Yew WP. Posterior tibialis tendon dysfunction: overview of evaluation and management. Orthopedics 2015;38:385 91. [10] Razeghi M, Batt ME. Foot type classification: a critical review of current methods. Gait Posture 2002;15:282 91. [11] Menz HB. Alternative techniques for the clinical assessment of foot pronation. J Am Podiatr Med Assoc 1998;88:119 29. [12] Kimura T, Thorhauer ED, Kindig MW, Sangeorzan BJ, Ledoux WR. Evaluation of the foot arch in partial weightbearing conditions. Foot Ankle Int 2022;43:113 22.

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[13] Shelton TJ, Singh S, Bent Robinson E, Nardo L, Escobedo E, Jackson L, et al. The influence of percentage weight-bearing on foot radiographs. Foot Ankle Spec 2019;12:363 9. [14] Sangeorzan BJ, Mosca V, Hansen ST. Effect of calcaneal lengthening on relationships among the hindfoot, midfoot, and forefoot. Foot Ankle 1993;14:136 41. [15] Saltzman CL, el-Khoury GY. The hindfoot alignment view. Foot Ankle Int 1995;16:572 6. [16] Ledoux WR, Rohr ES, Ching RP, Sangeorzan BJ. Effect of foot shape on the three-dimensional position of foot bones. J Orthop Res 2006;24:2176 86. [17] de Cesar Netto C, Schon LC, Thawait GK, da Fonseca LF, Chinanuvathana A, Zbijewski WB, et al. Flexible adult acquired flatfoot deformity: comparison between weight-bearing and non-weight-bearing measurements using cone-beam computed tomography. JBJS 2017;99:e98. [18] Shakoor D, de Cesar Netto C, Thawait GK, Ellis SJ, Richter M, Schon LC, et al. Weight-bearing radiographs and cone-beam computed tomography examinations in adult acquired flatfoot deformity. Foot Ankle Surg 2021;27:201 6. [19] Redmond AC, Crosbie J, Ouvrier RA. Development and validation of a novel rating system for scoring standing foot posture: the Foot Posture Index. Clin Biomech (Bristol, Avon) 2006;21:89 98. [20] Aquino MRC, Avelar BS, Silva PL, Ocarino JM, Resende RA. Reliability of Foot Posture Index individual and total scores for adults and older adults. Musculoskelet Sci Pract 2018;36:92 5. [21] Menz HB, Munteanu SE. Validity of 3 clinical techniques for the measurement of static foot posture in older people. J Orthop Sports Phys Ther 2005;35:479 86. [22] Vinicombe A, Raspovic A, Menz HB. Reliability of navicular displacement measurement as a clinical indicator of foot posture. J Am Podiatr Med Assoc 2001;91:262 8. [23] Cavanagh PR, Rodgers MM. The arch index: a useful measure from footprints. J Biomech 1987;20:547 51. 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[35] Sa´nchez-Rodrı´guez R, Martı´nez-Nova A, Escamilla-Martı´nez E, Pedrera-Zamorano JD. Can the Foot Posture Index or their individual criteria predict dynamic plantar pressures? Gait Posture 2012;36:591 5. [36] Buldt AK, Forghany S, Landorf KB, Levinger P, Murley GS, Menz HB. Foot posture is associated with plantar pressure during gait: a comparison of normal, planus and cavus feet. Gait Posture 2018;62:235 40. [37] Buldt AK, Allan JJ, Landorf KB, Menz HB. The relationship between foot posture and plantar pressure during walking in adults: a systematic review. Gait Posture 2018;62:56 67. [38] Buldt AK, Murley GS, Butterworth P, Levinger P, Menz HB, Landorf KB. The relationship between foot posture and lower limb kinematics during walking: a systematic review. Gait Posture 2013;38:363 72. [39] Hollander K, Zech A, Rahlf AL, Orendurff MS, Stebbins J, Heidt C. The relationship between static and dynamic foot posture and running biomechanics: a systematic review and meta-analysis. 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[44] Sadler S, Spink M, de Jonge XJ, Chuter V. An exploratory study investigating the effect of foot type and foot orthoses on gluteus medius muscle activity. BMC Musculoskelet Disord 2020;21:655. [45] Telfer S, Kindig MW, Sangeorzan BJ, Ledoux WR. Metatarsal shape and foot type: a geometric morphometric analysis. J Biomech Eng 2017;139. [46] Moore ES, Kindig MW, McKearney DA, Telfer S, Sangeorzan BJ, Ledoux WR. Hind- and midfoot bone morphology varies with foot type and sex. J Orthop Res 2019;37:744 59. [47] Peeters K, Schreuer J, Burg F, Behets C, Van Bouwel S, Dereymaeker G, et al. Alterated talar and navicular bone morphology is associated with pes planus deformity: a CT-scan study. J Orthop Res 2013;31:282 7. [48] Ilfeld FW. Pes planus: military significance and treatment with simple arch support. J Am Med Assoc 1944;124:281 3. [49] Garrow AP, Silman AJ, Macfarlane GJ. The Cheshire foot pain and disability survey: a population survey assessing prevalence and associations. Pain. 2004;110:378 84. [50] Mølgaard C, Lundbye-Christensen S, Simonsen O. High prevalence of foot problems in the Danish population: a survey of causes and associations. Foot 2010;20:7 11. [51] Menz HB, Dufour AB, Riskowski JL, Hillstrom HJ, Hannan MT. Association of planus foot posture and pronated foot function with foot pain: the Framingham foot study. Arthritis Care Res (Hoboken) 2013;65:1991 9. [52] Hagedorn TJ, Dufour AB, Riskowski JL, Hillstrom HJ, Menz HB, Casey VA, et al. Foot disorders, foot posture, and foot function: the Framingham foot study. PLoS One 2013;8:e74364. [53] Awale A, Hagedorn TJ, Dufour AB, Menz HB, Casey VA, Hannan MT. Foot function, foot pain, and falls in older adults: the framingham foot study. Gerontology 2017;63:318 24. [54] Menz HB, Dufour AB, Riskowski JL, Hillstrom HJ, Hannan MT. Foot posture, foot function and low back pain: the Framingham Foot Study. Rheumatol (Oxf) 2013;52:2275 82. [55] Levy JC, Mizel MS, Wilson LS, Fox W, McHale K, Taylor DC, et al. Incidence of foot and ankle injuries in West Point cadets with pes planus compared to the general cadet population. Foot Ankle Int 2006;27:1060 4. [56] Dixon S, Nunns M, House C, Rice H, Mostazir M, Stiles V, et al. Prospective study of biomechanical risk factors for second and third metatarsal stress fractures in military recruits. J Sci Med Sport 2019;22:135 9. [57] Kaufman KR, Brodine SK, Shaffer RA, Johnson CW, Cullison TR. The effect of foot structure and range of motion on musculoskeletal overuse injuries. Am J Sports Med 1999;27:585 93. [58] Yates B, White S. The incidence and risk factors in the development of medial tibial stress syndrome among naval recruits. Am J Sports Med 2004;32:772 80. [59] Ledoux WR, Shofer JB, Ahroni JH, Smith DG, Sangeorzan BJ, Boyko EJ. Biomechanical differences among pes cavus, neutrally aligned, and pes planus feet in subjects with diabetes. Foot Ankle Int 2003;24:845 50. [60] Ledoux WR, Shofer JB, Smith DG, Sullivan K, Hayes SG, Assal M, et al. Relationship between foot type, foot deformity, and ulcer occurrence in the high-risk diabetic foot. J Rehabil Res Dev 2005;42:665 72. [61] Levinger P, Menz HB, Morrow AD, Bartlett JR, Feller JA, Bergman NR. Relationship between foot function and medial knee joint loading in people with medial compartment knee osteoarthritis. J Foot Ankle Res 2013;6:33. [62] Telfer S, Abbott M, Steultjens MPM, Woodburn J. Dose-response effects of customised foot orthoses on lower limb kinematics and kinetics in pronated foot type. J Biomech 2013;46:1489 95. [63] Telfer S, Abbott M, Steultjens M, Rafferty D, Woodburn J. Dose-response effects of customised foot orthoses on lower limb muscle activity and plantar pressures in pronated foot type. Gait Posture 2013;38:443 9. [64] Unver B, Erdem EU, Akbas E. Effects of short-foot exercises on foot posture, pain, disability, and plantar pressure in pes planus. J Sport Rehabil 2020;29:436 40. [65] Okamura K, Fukuda K, Oki S, Ono T, Tanaka S, Kanai S. Effects of plantar intrinsic foot muscle strengthening exercise on static and dynamic foot kinematics: a pilot randomized controlled single-blind trial in individuals with pes planus. Gait Posture 2020;75:40 5. [66] Tang M, Wang L, You Y, Li J, Hu X. Effects of taping techniques on arch deformation in adults with pes planus: a meta-analysis. PLoS One 2021;16:e0253567. [67] Samson W, Van Hamme A, Sanchez S, Che`ze L, Jan SVS, Feipel V. Foot roll-over evaluation based on 3D dynamic foot scan. Gait Posture 2014;39:577 82. [68] Yan S, Li R, Shi B, Wang R, Yang L. Mixed factors affecting plantar pressures and center of pressure in obese children: obesity and flatfoot. Gait Posture 2020;80:7 13. [69] Kelly LA, Cresswell AG, Racinais S, Whiteley R, Lichtwark G. Intrinsic foot muscles have the capacity to control deformation of the longitudinal arch. J R Soc Interface 2014;11:20131188.

Chapter 29

Traumatic Foot and Ankle Injuries Scott Shawen1 and Tobin Eckel2 1

OrthoCarolina Foot and Ankle Institute, Charlotte, NC, United States, 2Walter Reed National Military Medical Center, Bethesda, MD, United States

Abstract Traumatic injuries to the foot and ankle are common in the general population. Often the result of falls or high-energy events such as car crashes, these injuries occur in all age groups and vary in severity and complexity, and can involve bone fractures and soft tissue injuries. This chapter covers the etiology, symptoms, diagnosis, and treatment of a range of traumatic injuries to the foot and ankle.

29.1

Introduction

Injuries to the foot and ankle resulting from a traumatic event are common across the general population. In many cases, these are the results of high-energy events such as car crashes, however in older individuals with lower bone strength they can result from lower energy events such as falls. These injuries vary in severity and complexity, and can involve bone fractures, soft tissue injuries, or a combination. This chapter covers the etiology, symptoms, diagnosis, and treatment of common traumatic injuries to the foot and ankle.

29.2

Pilon fractures

29.2.1 Etiology and pathophysiology Tibial plafond, or pilon fractures, generally occur by one of the two mechanisms: rotational forces or axial load. Rotational injuries tend to be lower energy injuries that typically have large intra-articular fracture fragments with minimal comminution and joint impaction, and less soft tissue injury. Axial loading injuries result from falls from height or motor vehicle accidents, and are usually higher energy injuries. These differ significantly from rotational injuries in that there is often marked articular and metaphyseal comminution, joint impaction, and severe soft tissue trauma, all of which result in a much less favorable prognosis [1].

29.2.2 Symptoms These injuries can occur in isolation but are often seen in polytraumatized patients. There may be visible bony deformity, but soft tissue damage will be most evident and can include swelling, ecchymosis, fracture blisters, skin tenting from underlying bony fragments, open wounds, and/or fractures. Patients should be monitored for the development of compartment syndrome [1,2].

29.2.3 Diagnostics/classification The initial work-up includes a vascular exam to confirm adequate perfusion and plain radiographs. A computed tomography (CT) scan is vital in characterizing the fracture and in developing the surgical plan; however, this is best done after initial reduction of the fracture. In fact, the CT scan can add information in 82% and change the initial surgical plan in 64% of patients [1,3]. Ru¨edi and Allgo¨wer initially classified these fractures based on articular displacement [4]: Type A being simple fractures with little or no articular displacement, Type B having articular displacement Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00010-X © 2023 Elsevier Inc. All rights reserved.

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without comminution, and Type C with both articular displacement and comminution. AO/OTA classification divides these injuries into three categories: extra-articular, partial articular, and complete articular. Each category is further divided based on the amount of comminution [5].

29.2.4 Treatment These fractures are almost universally managed surgically, although initial conservative measures may assist in improving outcomes. In the event of open wounds, intravenous antibiotics can be started initially, and gross contamination can be removed with saline irrigation performed to decrease foreign material and infectious burden. Any skin tenting or extruded bony fragments may be reduced to prevent further tension or compromise to the soft tissues. Medical optimization can also improve long-term results. Nicotine use and poorly controlled diabetes have both been shown to increase wound healing complications and the need for revision surgery, so initial management should include smoking cessation counseling and improved glycemic control [2]. Soft tissue impairment secondary to inflammation is greatest from 6 hours to 6 days post-injury, and this time frame should be avoided for definitive fixation of these fractures. Historically, when these injuries underwent immediate open reduction and internal fixation (ORIF), wound complications approached 100% [2,3]. While some experienced trauma surgeons still advocate for immediate ORIF, the vast majority favor a staged protocol. The goals of surgery are restoration of alignment, joint congruity, and stable fixation to allow for early motion [6]. The first stage focuses on adequate debridement, soft tissue stabilization, and temporary external fixation to maintain a provisional reduction (Fig. 29.1).

FIGURE 29.1 AP and lateral radiographs of intra-articular pilon fracture (A), (B); AP and lateral radiographs after external fixation. Note the improvement in length and alignment. (C), (D); CT scan after external fixation demonstrating a typical pattern with large posterior malleolar, medial, and anterior fragments (E); AP and lateral radiographs after posterolateral and anterolateral approaches to reduce and fix the fractures (F), (G).

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FIGURE 29.2 AP and lateral radiographs of intra-articular pilon fracture (A), (B); AP and lateral radiographs after closed reduction and splinting. This fracture did not undergo external fixation, as it was stable with splinting only (C), (D); CT scan demonstrating diffuse comminution (E); AP and lateral radiographs demonstrating indirect reduction techniques by fixing fibula through small distal incision, traction, and proximal fixation. Tibia was approached through an anterolateral incision and a plated construct used for reduction with dorsiflexion of the ankle reducing the posterior joint (F), (G).

This allows for soft tissue rest and a CT scan to plan for definitive ORIF when the soft tissue allows, which can be anywhere from 10 to 21 days post-injury (Fig. 29.2). There is no consensus on when to fix the fibula fracture. Potential advantages for fixing the fibula fracture in the initial stage is that it can assist in maintaining length and alignment of the tibia and may only require monoplanar external fixation. However, there is increased risk for fibular wound complications with initial plate fixation, and with significant fibular comminution, anatomic reduction may be difficult, and a malreduced fibula will impede tibial reduction at the time of definitive fixation. Further, placing a calcaneal transfixion pin to achieve biplanar stability is not technically demanding or more time consuming than a uniplanar external fixator [3]. The primary incision is usually determined after the CT scan is obtained, and therefore it may be wise to delay fibular fixation until the surgical approach or approaches have been determined, to insure the maximum possible skin bridge between incisions. Therefore if the location of the definitive incision is unknown, then the fibula should not be fixed at the initial stage of treatment [2]. In fact, a study by Tornetta et al. questioned whether the fibula needs to be fixed at all in pilon fractures. They found no difference in reduction or maintenance of alignment between those with or without fibula fixation, but they did demonstrate increased complications in the group treated with fibular fixation [6]. At the time of definitive fixation, there are several surgical approaches described, and these are based on the fracture pattern and existing soft tissue envelope. The posterolateral approach can be used for the fibula as well as the posterior tibia, and the anterior fragments can be visualized through an anteromedial or anterolateral approach. A straight anterior approach has the advantage of the best visibility of the anterior approaches and can also be used for revision to arthrodesis in the future [2]. In general, these fractures require a prolonged period of nonweight-bearing, typically between 8 12 weeks. It is generally agreed that range of motion exercises should be started early, but there is no clear consensus on what period

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of time defines early motion. We would recommend 6 weeks of cast immobilization, followed by initiation of motion and weightbearing at 12 weeks. Functional outcomes after pilon fracture tend to be poor, regardless of rehabilitation protocol [3].

29.3

Calcaneal fractures

29.3.1 Etiology and pathophysiology The calcaneus is the most commonly fractured tarsal bone, accounting for nearly 70% of all tarsal fractures. These fractures are typically the result of high-energy loading, commonly seen in falls from height or motor vehicle accidents. As the talus is driven down into the calcaneus, the primary fracture line occurs, separating the medial sustentaculum from the calcaneal tuberosity. Secondary fracture lines occur in the sagittal plane and are responsible for posterior facet comminution. This axial force results in the characteristic deformity of a shortened calcaneus and lateral displacement or blowout of the lateral wall. Calcaneal shortening leads to anterior ankle impingement and decreased push-off strength secondary to a shorter Achilles lever arm. This, in addition to lateral displacement, can lead to subfibular impingement, sural neuritis, and difficulty with shoe wear. Long-term sequelae include post-traumatic arthritis, pain, and stiffness. These injuries have major socioeconomic implications as the vast majority occur in young people and laborers [7].

29.3.2 Symptoms Patients present after high energy trauma, with pain, swelling, and inability to bear weight. Fracture blisters are also common. Open fractures occur in 1% 10% of calcaneal fractures. Posterior displaced tongue-type fractures are rare, but need to be recognized immediately, as the posterior skin tenting can lead to full thickness skin breakdown if not reduced urgently [8].

29.3.3 Diagnostics/classification Plain radiographs of the foot are obtained, with the lateral view demonstrating loss of height of the posterior facet. An axial view of the heel may demonstrate the amount of displacement and varus deformity. A CT scan is also routinely obtained to further delineate these injuries. The most commonly used classification system is the Sanders classification, which utilizes semi-coronal images through the posterior facet, and characterizes these fractures based on the number of posterior facet fracture fragments. The Essex-Lopresti classification scheme is based on plain films, and defines fractures as joint depression or tongue type, with the latter having a secondary fracture line that exits posteriorly, often with the Achilles attached to this tuberosity fragment, which leads to superior displacement and often tenting of the posterior skin [9].

29.3.4 Treatment Given the extensive soft tissue swelling and trauma, these fractures are typically placed in a well-padded splint with definitive treatment delayed until soft tissue swelling resolves. Exceptions include open fractures or those with skin tenting with impending open fractures; these require surgical debridement for existing open wounds or urgent reduction to prevent open wound formation. Foot compartment syndrome is also a concern with these injuries, although fasciotomies are rarely performed, as the main sequela of compartment syndrome is claw toe deformity. Others may advocate for early fixation for patients that are already admitted to the hospital for polytrauma. There is a trend toward early treatment (within the first week) with the sinus tarsi approach (Fig. 29.3). Nonoperative management is generally avoided, as it leads to malunion and higher incidence of subtalar arthritis. Malunion leads to anterior ankle impingement, subfibular impingement, pain, and difficulty with shoe wear [8,10]. Surgery can be delayed up to three weeks to allow the resolution of soft tissue swelling. However, with a lateral extensile approach, postsurgical wound breakdown can still occur in 25% of patients, with osteomyelitis occurring in 1% 4% of closed and up to 19% of open fractures [11]. Percutaneous techniques and limited sinus tarsi approaches have been utilized with fewer wound healing complications and earlier surgery, however, these are often limited to simpler fracture patterns with less articular comminution. Regardless of surgical approach, the goals are the same, including: restoring hindfoot alignment to prevent varus malunion, restoring posterior facet height and articular congruity, preventing anterior impingement and minimize secondary arthrosis, restoring calcaneal width, and preventing chronic subfibular impingement, sural neuritis, and difficulty with shoe wear. In higher energy fractures, primary subtalar

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FIGURE 29.3 Lateral radiograph and sagittal CT scan of calcaneus with minimal displacement (A), (B); Lateral and axial radiographs, and sagittal CT scan of same calcaneus treated within one week of injury with a sinus tarsi approach (C), (D), and (E).

arthrodesis can be performed after reduction of the fracture (Fig. 29.4). Another advantage of more anatomic restoration with surgery includes fewer complications and better outcomes with delayed subtalar arthrodesis when compared to arthrodesis after nonoperative management [10,12].

29.4

Talus fractures

29.4.1 Etiology and pathophysiology Talus fractures are the second most common tarsal fracture (those affecting the calcaneus are the most common) and make up about 0.15% of all fractures [13]. As with most tarsal fractures, they occur from high-energy trauma but can be a result of sports injuries as well. Medial sided forefoot compression causes talar head fractures, forced dorsiflexion of the foot is the primary cause of talar neck fractures, and axial compression between the tibia and calcaneus is the primary cause of talar body fractures [13,14].

29.4.2 Symptoms After high-energy trauma, there will be significant deformity and/or swelling in the ankle and hindfoot, particularly with displaced talar neck fractures, which are often associated with dislocation. Clinicians should be suspicious of open injuries or pending open injuries with associated dislocations. With lower energy trauma (sports, snowboarding) clinicians should be suspicious of an occult talus fracture with negative radiographs but swelling and inability to weight bear.

29.4.3 Diagnostics/classification Ankle and foot radiographs should be taken for all patients with suspected foot or ankle fractures. Any irregularities of the talus seen on radiographs should be further evaluated with a CT scan [13,14]. Surgeons should consider contralateral radiographs for comparison.

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FIGURE 29.4 Lateral radiograph and axial CT scan demonstrating a high-grade calcaneal fracture (Sanders grade 4) (A), (B); Lateral and axial radiographs after open reduction through a sinus tarsi approach and primary subtalar arthrodesis (C), (D). Credit: Images supplied by the authors.

Talus fractures are first classified by anatomic location: head (10%), neck (50%) and body (38%). Neck fractures are further described as either nondisplaced (OTA 81-B1; Hawkins grade 1), subtalar joint incongruity (81-B2; Hawkins grade 2), subtalar and tibiotalar joint incongruity (81-B3; Hawkins grade 3), or complete peritalar dislocation (Hawkins grade 4). Body fractures (posterior to lateral tubercle) are further described as talar dome (OTA 81-C1), with subtalar involvement (B1 C2), or subtalar and tibiotalar joint involvement (81-C3) [14].

29.4.4 Treatment Because the tenuous extraosseous blood supply of the talus, there are high rates of osteonecrosis after fracture displacement. Higher rates occur with more displacement, particularly seen with neck fractures. Therefore, nonoperative management is reserved for nondisplaced fractures only. Immobilization with protected weightbearing for 6 8 weeks is typical for nondisplaced fractures. Small avulsions injuries are treated symptomatically and usually do not require immobilization [13,14]. Displaced fractures are best treated with ORIF (Fig. 29.5). External fixation may be needed for temporary fixation in the setting of poor soft tissue envelope and instability. Fracture dislocations and open injuries are treated emergently and require open reduction. Internal and/or external fixation is used at time of open reduction.

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FIGURE 29.5 AP and lateral ankle radiographs of talar body fracture with subtalar dislocation (A), (B); Intraoperative fluoroscopy of talar fixation after reduction. Dual incisions with multiple k-wires were used to reduce fractures until fixation was completed (C), (D), and (E). Credit: Images supplied by the authors.

For closed injuries, accuracy of reduction is more important than timing of fixation. Emergent reduction of talar neck fractures is no longer considered important in lessening rates of osteonecrosis. Neck fractures are best treated with dual approach, medial and lateral, and often require medial and lateral fixation. Care must be taken not to compress comminution of the medial side, as this can lead to varus malunion. Talar head fractures are approached medially and treated with extra-articular screws and/or intra-articular countersunk or headless screws. Medial column length preservation is needed and may require bridge plating or external fixation. Body fractures are treated primarily with screws, countersunk or headless if overlying articular cartilage (60% of the talus is intra-articular). Screws may be inserted from the anteromedial, anterolateral, posterolateral, or posteromedial direction or through a medial malleolus osteotomy. Small lateral process fractures are excised or fixed with a screw depending on size. Osteochondral fragments are fixed with headless screws or bioabsorbable pins [13 15].

29.5

Tarsometatarsal (Lisfranc) injuries

29.5.1 Etiology and pathophysiology Ligamentous ruptures and/or fractures across the tarsometatarsal (TMT) joints are commonly referred to as “Lisfranc” injuries. The primary ligament involved connects the medial cuneiform to the base of the second metatarsal but injuries can occur across all 5 TMT joints and the strong plantar ligaments connecting metatarsals 2 5. Over 20% of TMT

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injuries are missed. To understand etiology and pathophysiology, it is important to distinguish lower energy injuries (sports, twisting/falls) from fracture/dislocations that typically occur with high-energy injuries (motor vehicle accidents, falls from height, crush injuries) [16,17]. The true TMT fracture dislocation is a rare injury that occurs 1 in 50,000 people in the United States or 0.2% of all fractures [16,18]. These injuries occur from forceful torsion of the foot, axial loads in equinus, or direct crush injuries onto the midfoot. Midfoot sprains, as classically diagnosed with lower-energy injuries, are the second most common sports foot injury, and make up 16% of all sports injuries [18]. Lower energy injuries tend to occur from indirect mechanisms including axial load to a plantarflexed foot, or a twisting force through the midfoot particularly seen with cleated athletes [16].

29.5.2 Symptoms High-energy injuries will have significant midfoot swelling and ecchymosis (plantar and dorsal) at the midfoot, possible deformity, soft tissue injury as well as possible foot compartment syndrome. The lower-energy injuries may or may not have swelling in the midfoot, inability to bear weight, tenderness over the midfoot as well as plantar ecchymosis. The missed TMT fracture dislocation injury often presents with continued pain with weightbearing after several months with or without previous immobilization [16,18].

29.5.3 Diagnostics/classification For high-energy injuries, AP, lateral and oblique radiographs of the foot along with a CT scan to evaluate the fracture pattern are needed. For the lower-energy injury, in addition to radiographs of the foot, bilateral weightbearing AP radiographs are needed to compare to the contralateral side and evaluate for gapping at the first intermetatarsal space at the base as well as subluxation of the second and third TMT joints. A CT scan can help to evaluate fractures and joint incongruity and MRI can evaluate the continuity of the medial cuneiform second metatarsal ligament in equivocal cases or for patients that cannot bear weight [16 18]. Abduction stress radiographs or fluoroscopy can demonstrate instability as well [19]. This should be done after anesthetizing the patient or in the operating room under anesthesia for surgical planning. High-energy injuries are often described using the Myerson classification. Type A have total incongruity (lateral or medial) of all five TMT joints. Type B have partial incongruity (B1-medial, B2-partial lateral, B3-total lateral). Type C is divergent (C1-partial, C2-total) [18]. Low-energy injuries are best described by Nunley and Vertullo with Stage 1 having no diastasis at the medial cuneiform second metatarsal space with weightbearing X-rays, Stage 2 having 2 5 mm of diastasis of AP weightbearing X-rays and Stage 3 having greater than 5 mm diastasis and loss of longitudinal arch seen on lateral foot weightbearing X-rays [16,17].

29.5.4 Treatment Only Stage 1 injuries should be treated without surgery. Protected weightbearing in an orthosis or a short leg cast is an option for 4 6 weeks, but many patients do fine with orthosis and weightbearing as tolerated with avoidance of prolonged standing or walking for several weeks. Almost all other injuries do better long term with surgery. The key to surgery is anatomic reduction and options consist of primary arthrodesis or open reduction internal fixation (Figs. 29.6 and 29.7). Controversy exists on choosing arthrodesis vs internal fixation. While similar outcomes have been reported with both techniques, there is a higher rate of secondary hardware removal and the development of post-traumatic arthritis in 20% 50% of patients undergoing ORIF, which may necessitate salvage arthrodesis [20]. In an athletic population, there is a trend toward higher return to play with primary arthrodesis compared to open reduction internal fixation. Additionally, primary arthrodesis demonstrates better outcomes compared to salvage arthrodesis after failed internal fixation [20,21].

29.6

Metatarsal fractures

29.6.1 Etiology and pathophysiology Metatarsal fractures are relatively common, accounting for 6% of all fractures seen in the primary care setting. These fractures occur through direct or indirect forces. Direct force can happen acutely, such as a crushing type injury from heavy objects landing on the foot, or it can be more chronic in nature, for instance, repetitive microtrauma leading to stress fractures. Twisting injuries in sports is an example of an indirect force that can lead to metatarsal fracture. The

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FIGURE 29.6 AP, oblique, and lateral radiographs of TMT fracture and dislocation involving all five joints with lateral shift of metatarsals (A) (C); AP, oblique, and lateral radiographs after open reduction and primary fusion of joints 1 3, pinning of joints 4 and 5 (D) (F). Credit: Images supplied by the authors.

fifth metatarsal is the most commonly fractured metatarsal, and can be the result of an acute, twisting injury, or a chronic stress fracture [22,23].

29.6.2 Symptoms Similar to any other fracture, patients often present with pain overlying the metatarsal, swelling, and ecchymosis, and difficulty or inability to bear weight. For stress fractures, they often report an insidious onset of pain that gradually worsens and is often exacerbated by increased impact activity. These may often coincide with a recent increase in baseline level of activity, such as training for a marathon, or may be more subtle, like wearing a less supportive pair of shoes.

29.6.3 Diagnostics/classification Plain radiographs are sufficient in the evaluation of metatarsal fractures, with the one exception being metatarsal base fractures, where CT or MRI may be warranted if there is an expected midfoot fracture/dislocation or associated midfoot ligamentous injury. Metatarsal fractures are often classified anatomically as neck, shaft, or base fractures [22]. Fifth metatarsal fractures are often classified by anatomic zones. Zone 1 represents a proximal tuberosity fracture, the result of an avulsion force by either the lateral band of the plantar fascia or the insertion of the peroneus brevis and occurs with an acute inversion injury. Zone 2 fractures, commonly known as Jones fractures, occur in a watershed area of the bone, 1.5 2 cm distal to the tuberosity with the fracture exiting into the 4 5 meta-diaphyseal junction. Zone 3 fractures occur in the metatarsal shaft. Torg et al. also classified these fractures based on acuity and thus healing potential, with type I representing an acute injury, type II representing a delayed union where there was a history of previous injury but incomplete radiographic healing and persistent pain, and type III representing a nonunion, with chronic pain and no radiographic healing. There is also a characteristic spiral oblique fracture of the mid to distal shaft of the fifth metatarsal fracture, commonly referred to as the “dancer’s” fracture, as it was historically reported in ballet dancers [23 25].

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FIGURE 29.7 Bilateral standing radiographs of left TMT injury. Note the diastasis at the base of the second metatarsal which extends into the intercuneiform joint and impaction fracture of the navicular (A), (B); MRI of the left foot demonstrating disruption of the medial cuneiform second metatarsal ligament at the base of the second metatarsal (C); After open reduction and screw fixation with elevation of the impaction of the navicular with bone grafting (D); follow up 9 months later after removal of screws at 6 months from injury (E). Credit: Images supplied by the authors.

29.6.4 Treatment There are no clear guidelines for when to operate on metatarsal fractures, but these injuries are often managed nonoperatively with protected weightbearing in a controlled ankle movement boot or post-op shoe. Fifth metatarsal fractures, on the other hand, typically require 6 8 weeks of non weightbearing when managed nonoperatively. Generally, up to 4 mm of translation and shortening and up to 10 degrees of angulation is well tolerated in metatarsal shaft fractures. The remaining intact metatarsals serve as internal splints and the ligamentous connections confer stability and help to minimize fracture displacement. Neck fractures are usually treated nonoperatively as well, unless there is significant rotational deformity or plantar angulation that could lead to metatarsalgia [22]. Surgical options for metatarsal neck and shaft fractures include closed reduction and percutaneous pinning, open reduction and percutaneous pinning, and open reduction internal fixation (Fig. 29.8). Metatarsal base fractures in the absence of any midfoot dislocation or ligamentous injury, are normally managed nonoperatively [22]. However, nonunion can be managed operatively with either open reduction and fixation for larger fragments (Fig. 29.9), or excision for smaller fragments. Treatment of fifth metatarsal fractures is based on location and chronicity. Zone 1 fractures are almost universally managed nonoperatively. Acute zone 2 and 3 fractures can also be managed nonoperatively, although there is a trend toward surgery for zone 2 fractures to promote earlier fracture healing and return to sports (Fig. 29.10). Stress fractures, delayed unions, and nonunions should undergo surgery, and are typically augmented with bone graft [23,26,27]. Intramedullary screw fixation has superior bending stiffness and resistance to torsional loading when compared to plate fixation; thus, we routinely use a solid intramedullary screw for fixation. It is also important to assess the foot for any

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FIGURE 29.8 AP radiograph demonstrating multiple displaced metatarsal neck fractures and dislocated fifth metatarsophalangeal joint (A); AP radiograph after open reduction and internal fixation with k-wires (B); AP radiograph 4 months after injury. K-wires were removed at 6 weeks (C). Credit: Images supplied by the authors.

FIGURE 29.9 AP and lateral radiographs of non-united fifth metatarsal base fracture (Zone I) at 9 months from injury (A), (B); AP and lateral radiographs after open reduction and internal fixation with calcaneal autograft. Patient offered excision of the fragment given the size, but preferred fixation (C), (D). Credit: Images supplied by the authors.

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FIGURE 29.10 Lateral, oblique, and AP radiographs of fifth metatarsal “Jones” or Zone II fracture (A) (C); lateral, oblique, and AP radiographs after intramedullary screw fixation (D) (F). Credit: Images supplied by the authors.

cavovarus deformity, which can lead to stress fractures of the fifth metatarsal, and increased nonunion if not addressed at the time of surgery [23,28].

29.7

Midfoot crush injuries

29.7.1 Etiology and pathophysiology Crush injuries are unique in that they can encompass a combination of any fractures and dislocations about the foot. They typically result from high-energy axial load injuries and have significant soft tissue damage. This leads to an increased incidence of compartment syndrome and the need for soft tissue coverage [29].

29.7.2 Symptoms These injuries often occur in the setting of polytrauma. Diffuse swelling, open fractures, and extensive soft tissue injury are common findings upon presentation [29].

29.7.3 Treatment Foot fasciotomies may increase the risk of wound infection and need for soft tissue coverage. Claw toe deformity is the most common sequela of compartment syndrome in the foot, secondary to the extrinsic musculature overpowering the weak intrinsic muscles. Chronic pain remains a common finding in high-energy foot trauma, with no evidence to suggest a preventative benefit with fasciotomies at the time of injury. For these reasons, foot fasciotomies are rarely performed [30]. Because of the lack of rotational flaps to cover the foot, free tissue flaps are the main option for coverage. It is critical to take a multi-disciplinary approach to include plastic surgery early on in the management of these injuries [31].

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29.8

473

Acute ankle sprains

29.8.1 Etiology and pathophysiology Ankle sprains are among the most common athletic injuries, accounting for nearly half of all sport related injuries [32,33]. The incidence rates of ankle sprains in the general population range between 5 7 per 1000 person-years, although it is much higher in more athletic populations [34,35]. These injuries are the result of ankle inversion, with injury most common in the anterior talofibular ligament, followed by the calcaneofibular ligament, and lastly the posterior talofibular ligament. While less common, the deltoid ligament may also be injured, which often leads to longer recovery times [36].

29.8.2 Symptoms Ankle sprains typically present with lateral swelling and pain with ambulation. Tenderness is often localized to the lateral ligament complex, and ecchymosis is often present. Roughly 95% of ankle sprains will resume their previous level of sports activity within 6 weeks of injury. However, nearly half will still have some residual pain at 6 months postinjury [37].

29.8.3 Treatment Acute ankle sprains are almost always managed by nonoperative methods. Functional ankle rehabilitation results in quicker return to work, mobility, and preinjury activity level when compared to acute surgical repair/reconstruction [36]. Bracing may be used initially during rehabilitation, but there is little evidence to suggest any protective benefit from prophylactic bracing [38].

29.9

Syndesmosis tears

29.9.1 Etiology and pathophysiology Syndesmotic injuries, commonly referred to as high ankle sprains, account for roughly 10% of all ankle sprains, but perhaps as many as 25% of ankle sprains in collision sports. The mechanism of injury is believed to be external rotation while the foot is in a dorsiflexed position. These injuries have a longer recovery time and higher likelihood of long-term dysfunction when compared to low ankle sprains. These injuries rarely occur in isolation, and are often accompanied by a deltoid ligament injury or an ankle fracture [39]. The role of syndesmosis injury in ankle arthritis is controversial, currently with no clear evidence that a high ankle sprain leads to an increased incidence of traumatic ankle arthritis [40].

29.9.2 Symptoms These injuries may present similarly to a low ankle sprain, except the tenderness is localized to the anterior inferior tibiofibular ligament or the interosseous membrane. Proximal tibia-fibular compression that reproduces distal tibiafibular pain, known as the squeeze test, is indicative of syndesmotic injury. Similarly, an external rotation stress test that reproduces pain over the syndesmosis is also suggestive of syndesmotic injury. The Cotton test, which depicts lateral or posterior fibula translation relative to the tibia, is typically done intraoperatively while fixing a concomitant ankle fracture [39].

29.9.3 Treatment Purely ligamentous injuries are treated nonoperatively, although patients need to be counseled that recovery is longer than the more common inversion ankle injury. Syndesmotic injuries with associated fracture and/or deltoid ligament injury typically require operative fixation, although techniques vary widely. Screws and suture button constructs are the most common fixation methods with or without deltoid ligament repair (Fig. 29.11). No difference has been shown when using 3.5 or 4.5 mm screws, or when achieving three vs four cortical fixations. If a posterior malleolar fracture is present, some advocate fixing this fracture to stabilize the syndesmosis [39]. Regardless of fixation construct, post-reduction CT scans after screw fixation demonstrate malreduction of the syndesmosis between 36% 52% of the time [41]. Furthermore, 89% spontaneously reduce after screw removal. For these

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FIGURE 29.11 AP radiograph of the ankle demonstrating syndesmotic injury with deltoid ligament involvement and fibula fracture (A); AP and lateral radiographs after arthroscopy (medial cartilage defect debrided), open reduction and internal fixation of the fibula fracture, syndesmotic fixation with suture button devices, and repair of the deltoid ligament (B), (C). Credit: Images supplied by the authors.

reasons we recommend routine screw removal or the use of suture button constructs for these injuries, assuming that the fibula has been fixed if fractured and is length stable.

29.10 Achilles tendon rupture 29.10.1 Etiology and pathophysiology Sporting activity accounts for nearly 70% of all Achilles tendon ruptures, with basketball being the most common [42]. The increase in Achilles tendon ruptures has been credited to the expanding number of older individuals participating in high-impact sports. This is likely multifactorial, with normal tendon degeneration over time, the high forces seen across the Achilles tendon with running and jumping that can exceed 12 times the body weight, and the relative vascularity to the midsubstance of the Achilles [43].

29.10.2 Symptoms Patients typically describe an audible pop or the sensation that someone kicked them in the back of the leg. Pain is variable, but difficulty with weight-bearing and lack of push-off strength are common. Ecchymosis and swelling are typical and often there is a visual and/or palpable defect in the tendon. Lack of plantarflexion with calf squeeze (Thompson test) and asymmetric resting equinus posture when prone (Matles test) are both pathognomonic for an Achilles tendon rupture [44]. Rupture of the tendon can be readily visualized using MRI (Fig. 29.12).

29.10.3 Treatment Recent literature has demonstrated equivalent outcomes with operative and non-operative treatment for acute Achilles tendon rupture when using a functional rehabilitation protocol. This has led to a greater trend toward non-operative management for these injuries [45]. However, some patient populations do benefit from surgical repair. Athletic populations return to previous level of sport sooner and more often than those treated nonoperatively [44]. And while rerupture rates are equivalent across all demographics, males under the age of 40 have higher rerupture rates with nonoperative treatment when compared to females and males over 40 [46]. Non-operative treatment negates the risk of wound complications and infection, but these risks range between 1% 4% with operative treatment, and are particularly low with minimally invasive techniques [46,47]. Overall, it is reasonable based on the current evidence, to treat acute Achilles tendon ruptures either with or without surgery.

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FIGURE 29.12 Sagittal MRI of the ankle demonstrating Achilles tendon rupture. Credit: Images supplied by authors.

29.11 Areas of future research Biomechanical studies in the field of traumatic injuries to the foot and ankle remain important avenues of research. Many questions remain, particularly in the area of fracture healing where understanding how to improve surgical techniques and rehabilitation care to reduce non-union rates and other complications, and addressing these is a common goal in orthopedics.

References [1] Barei DP, Nork SE. Fractures of the tibial plafond. Foot Ankle Clin 2008;13(4):571 91. [2] Liporace FA, Yoon RS. Decisions and staging leading to definitive open management of pilon fractures: where have we come from and where are we now? J Orthop Trauma 2012;26(8):488 98. [3] Crist BD, et al. Pilon fractures: advances in surgical management. J Am Acad Orthop Surg 2011;19(10):612 22. [4] Ru¨edi TP, Allgo¨wer M. The operative treatment of intra-articular fractures of the lower end of the tibia. Clin Orthop Relat Res 1979;138:105 10. PMID 376196. [5] Borrelli Jr. J, Ellis E. Pilon fractures: assessment and treatment. Orthop Clin North Am 2002;33(1):231 45 x. [6] Kurylo JC, et al. Does the fibula need to be fixed in complex pilon fractures? J Orthop Trauma 2015;29(9):424 7. [7] Jackson 3rd JB, et al. Distraction subtalar arthrodesis. Foot Ankle Clin 2015;20(2):335 51. [8] Swords MP, Penny P. Early fixation of calcaneus fractures. Foot Ankle Clin 2017;22(1):93 104. [9] Epstein N, Chandran S, Chou L. Current concepts review: intra-articular fractures of the calcaneus. Foot Ankle Int 2012;33(1):79 86. [10] Kiewiet NJ, Sangeorzan BJ. Calcaneal fracture management: extensile lateral approach vs small incision technique. Foot Ankle Clin 2017;22(1):77 91. [11] Clare MP, Crawford WS. Managing complications of calcaneus fractures. Foot Ankle Clin 2017;22(1):105 16. [12] Radnay CS, Clare MP, Sanders RW. Subtalar fusion after displaced intra-articular calcaneal fractures: does initial operative treatment matter? J Bone Jt Surg Am 2009;91(3):541 6. [13] Fortin PT, Balazsy JE. Talus fractures: evaluation and treatment. J Am Acad Orthop Surg 2001;9(2):114 27. [14] Grear BJ. Review of talus fractures and surgical timing. Orthop Clin North Am 2016;47(3):625 37. [15] Early JS. Talus fracture management. Foot Ankle Clin 2008;13(4):635 57. [16] Lewis Jr. JS, Anderson RB. Lisfranc injuries in the athlete. Foot Ankle Int 2016;37(12):1374 80. [17] Coetzee JC. Making sense of lisfranc injuries. Foot Ankle Clin 2008;13(4):695 704 ix. [18] Eleftheriou KI, Rosenfeld PF. Lisfranc injury in the athlete: evidence supporting management from sprain to fracture dislocation. Foot Ankle Clin 2013;18(2):219 36.

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[19] Coss HS, et al. Abduction stress and AP weightbearing radiography of purely ligamentous injury in the tarsometatarsal joint. Foot Ankle Int 1998;19(8):537 41. [20] Hawkinson MP, et al. Outcomes of Lisfranc Injuries in an Active Duty Military Population. Foot Ankle Int 2017;38(10):1115 19. [21] MacMahon A, et al. Return to sports and physical activities after primary partial arthrodesis for lisfranc injuries in young patients. Foot Ankle Int 2016;37(4):355 62. [22] Armagan OE, Shereff MJ. Injuries to the toes and metatarsals. Orthop Clin North Am 2001;32(1):1 10. [23] Bowes J, Buckley R. Fifth metatarsal fractures and current treatment. World J Orthop 2016;7(12):793 800. [24] Fetzer GB, Wright RW. Metatarsal shaft fractures and fractures of the proximal fifth metatarsal. Clin Sports Med 2006;25(1):139 50 x. [25] Aynardi M, et al. Outcome of nonoperative management of displaced oblique spiral fractures of the fifth metatarsal shaft. Foot Ankle Int 2013;34(12):1619 23. [26] O’Malley M, et al. Operative treatment of fifth metatarsal jones fractures (Zones II and III) in the NBA. Foot Ankle Int 2016;37(5):488 500. [27] Solan M, Davies M. Nonunion of fifth metatarsal fractures. Foot Ankle Clin 2014;19(3):499 519. [28] Huh J, et al. Biomechanical comparison of intramedullary screw vs low-profile plate fixation of a jones fracture. Foot Ankle Int 2016;37 (4):411 18. [29] DiDomenico LA, Thomas ZM. Midfoot crush injuries. Clin Podiatr Med Surg 2014;31(4):493 508. [30] Dodd A, Le I. Foot compartment syndrome: diagnosis and management. J Am Acad Orthop Surg 2013;21(11):657 64. [31] Heitmann C, Levin LS. The orthoplastic approach for management of the severely traumatized foot and ankle. J Trauma 2003;54(2):379 90. [32] Bastien M, et al. Alteration in global motor strategy following lateral ankle sprain. BMC Musculoskelet Disord 2014;15:436. [33] Chan KW, Ding BC, Mroczek KJ. Acute and chronic lateral ankle instability in the athlete. Bull NYU Hosp Jt Dis 2011;69(1):17 26. [34] Bulathsinhala L, et al. Epidemiology of ankle sprains and the risk of separation from service in U.S. army soldiers. J Orthop Sports Phys Ther 2015;45(6):477 84. [35] Waterman BR, et al. Epidemiology of ankle sprain at the United States Military Academy. Am J Sports Med 2010;38(4):797 803. [36] Al-Mohrej OA, Al-Kenani NS. Acute ankle sprain: conservative or surgical approach? EFORT Open Rev 2016;1(2):34 44. [37] Cameron KL, Owens BD, DeBerardino TM. Incidence of ankle sprains among active-duty members of the United States Armed Services from 1998 through 2006. J Athl Train 2010;45(1):29 38. [38] Newman TM, Gay MR, Buckley WE. Prophylactic ankle bracing in military settings: a review of the literature. Mil Med 2017;182(3): e1596 602. [39] Hunt KJ, et al. High ankle sprains and syndesmotic injuries in athletes. J Am Acad Orthop Surg 2015;23(11):661 73. [40] Kortekangas T, et al. Effect of syndesmosis injury in SER IV (Weber B)-type ankle fractures on function and incidence of osteoarthritis. Foot Ankle Int 2015;36(2):180 7. [41] Song DJ, et al. The effect of syndesmosis screw removal on the reduction of the distal tibiofibular joint: a prospective radiographic study. Foot Ankle Int 2014;35(6):543 8. [42] Raikin SM, Garras DN, Krapchev PV. Achilles tendon injuries in a United States population. Foot Ankle Int 2013;34(4):475 80. [43] Burrus MT, et al. Achilles tendon repair in obese patients is associated with increased complication rates. Foot Ankle Spec 2016;9(3):208 14. [44] Gross CE, Nunley 2nd JA. Acute achilles tendon ruptures. Foot Ankle Int 2016;37(2):233 9. [45] Renninger CH, et al. Operative and nonoperative management of achilles tendon ruptures in active duty military population. Foot Ankle Int 2016;37(3):269 73. [46] Cooper MT. Acute achilles tendon ruptures: does surgery offer superior results (and other confusing issues)? Clin Sports Med 2015;34 (4):595 606. [47] Hsu AR, et al. Clinical outcomes and complications of percutaneous achilles repair system vs open technique for acute achilles tendon ruptures. Foot Ankle Int 2015;36(11):1279 86.

Chapter 30

The Pediatric Foot Julie Stebbins1,2 and Max Mifsud1 1

Oxford University Hospitals NHS Foundation Trust, Oxford, United Kingdom, 2Nuffield Department of Orthopaedics, Rheumatology and

Musculoskeletal Sciences (NDORMS), Oxford University, Oxford, United Kingdom

Abstract The child’s foot is not just a smaller version of an adult foot. It differs in terms of structure and function and in biomechanical properties such as the elastic modulus of its bone, ligaments, and tendons. The foot is vital to activities of everyday living; therefore, any compromise in foot function is likely to adversely affect the overall quality of life. Ensuring adequate function of the foot is especially important in children as normal growth and development of the foot are dependent on experiencing appropriate loading and muscle forces. There are many different pathologies that can interfere with the normal development of children’s feet and these range from the idiopathic, such as flexible flat feet, to those associated with other conditions such as cerebral palsy and Charcot-MarieTooth disease. To fully understand the anatomical abnormalities, and subsequent effects on function, appropriate biomechanical assessments are required. These need to be tailored to pediatric feet. For example, their smaller foot size limits and sometimes precludes certain investigations. Currently, there is a lack of research documenting the long-term effects of interventions to accommodate or correct foot deformities through the use of orthoses or surgery. Research that provides this evidence would be valuable in improving our understanding of the effect of specific interventions, and how the pediatric foot responds to variations in intrinsic and extrinsic forces applied to it over time. This chapter explores the biomechanics of the pediatric foot in relation to common foot pathologies.

30.1

Introduction

The child’s foot is not just a smaller version of an adult foot. It differs in terms of structure and function and in biomechanical properties, largely due to the relatively lower modulus of elasticity in pediatric bony tissue compared to soft tissue [1]. The ratio of soft to bony tissue, position and orientation of the bones, as well as the amount and composition of soft tissue within the foot all vary as a child’s foot develops and matures. The foot is proportionally at its longest compared to body height at the time of birth. By 3 years of age, the foot has already achieved two thirds of its final length. It continues to grow rapidly (at around two millimeters per month) until the age of 3 years and then slows to around one millimeter per month between the ages of 3 and 5 years [2]. After this, growth is further slowed throughout the rest of childhood. The foot is normally the first part of the body to complete the growth process, achieving its full length around 3 years prior to skeletal maturity [3]. At the time of birth, the foot is made up of a relatively high proportion of cartilage and soft tissue. Progressive ossification of the tarsal bones begins during the embryonic period and continues to around 10 years of age [4]. The physeal growth plates of the metatarsals can take longer to full ossify (Fig. 30.1). Typically, the distal and proximal phalanges, talus and calcaneus are the first to begin the ossification process. The bones of the midfoot are generally the last to completely ossify. The center of ossification for the navicular may not emerge until 3 years of age, although this is highly variable [4,5]. The orientation of the bones of the foot also changes throughout growth. The infant talus is rotated and positioned more medially than in an adult foot [6]. As standing and walking ability is gained, the calcaneus progresses from an everted to more neutral position [7]. The positions of the malleoli also rotate externally as a baby progresses from crawling to walking and this accounts for the development of the normal foot progression angle [8]. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00030-5 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 30.1 Timing of ossification centers of the foot. Image by macrovector on Freepik.

It is commonly accepted that children are born with flat feet and that the medial, longitudinal arch develops over time [9]. The age at which the arch fully develops remains controversial, however. A recent review that included 34 cross-sectional studies of foot posture data on children aged 10 months to 18 years concluded that no definitive age could be specified at which the foot posture of children ceases to develop, primarily due to differing techniques used to quantify arch height [7]. This is further complicated by the fact that gender and body composition also influence development of the arch [10]. In addition to changing skeletal morphology, the soft tissue structures of the pediatric foot also develop over time. The collagen fibers within tendons and ligaments have fewer cross-links in young children, meaning they are more elastic compared to adult soft tissues. In addition, muscles grow in size and force production capability as the foot is exposed to increasing loads and activity. Finally, the fat pads under the heel, toes, and arch of the foot continue to change and develop, particularly as the child learns to walk [10]. The fat pad under the medial arch is gradually reduced between the ages of two and six [10]. In the infant foot, the heel and toe pads are thicker and the heel pad is located more distally than in the adult foot [11]. Differences in the structure and composition of children’s feet compared to adults’ have clear implications on gait and function. For example, the relatively large foot size of children, compared to body weight, contributes to lower peak plantar pressures during walking [12]. The increased flexibility of the pediatric foot makes it less susceptible to certain ligamentous injuries, but also means that growth and development can be influenced by both extrinsic (such as shoe wear) and intrinsic (such as imbalanced muscle strength across a joint) factors. The variable structure of the medial longitudinal arch has potential effects on the development and function of the more proximal joints of the lower limbs [13]. For example, a flattened arch may lead to altered coronal plane knee moments, which could contribute to the development of knee pain. It is clear, therefore, that children’s feet are not just smaller versions of adult feet. The differences in structure, composition, and proportions of children’s feet compared to adult feet have implications for how they are assessed, and how to interpret deviations from “normal” values.

30.2

Common pathologies affecting pediatric feet

Deviations in structure and/or function of the foot during growth primarily affect the child’s ability to perform weightbearing activities. This in turn, has further deleterious effects on the development of the child’s foot, as appropriate loading and muscle forces are required for maturation of both bony and soft tissue structures. Common foot deformities that impair biomechanical function of the foot and inhibit normal growth in children are outlined in the following sections.

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30.2.1 Congenital foot deformities Clubfoot or congenital talipes equinovarus is a condition that develops within the first trimester of pregnancy and is generally diagnosed antenatally on ultrasound, or clinically at birth. It is characterized by the presence of hindfoot equinus and varus with a cavus midfoot and internally rotated forefoot (Fig. 30.2). It is more common in males than females and can be unilateral or bilateral [14]. It is generally agreed to be multifactorial in etiology [15]. Previously, children diagnosed with clubfoot were treated surgically. However, the Ponseti technique of serial casting became the most accepted form of treatment worldwide from the early 2000s [14]. It is performed during the first few months of life. Using the biomechanical principle of creep (gradual deformation of a structure under constant load) sequential manipulations of the foot to correct the various deviations are performed and the corrected position held with long leg casting for several weeks per each manipulation before progressing to the next. At the end of the serial manipulations and casting, 80% of children will require an Achilles tendon tenotomy to correct the final hindfoot equinus. Following this, the child is placed in foot and ankle orthoses for varying amounts of time during the day and night for the first 4 years of life. Dynamic forefoot varus can sometimes recur in children previously treated with the Ponseti method. If appropriate gait analysis confirms this as an isolated aberration, this is often treated successfully with a tibialis anterior tendon transfer. In the Western world, it is rare to encounter untreated clubfoot beyond infancy, although in developing countries this is still a relatively common presentation. There is some evidence in the literature that Ponseti casting may also be effective in treating the older child with neglected clubfoot [16]. While there is general agreement in the literature that Ponseti casting provides superior results to surgical correction, the reported success rate still varies considerably [17]. Even if the initial treatment of clubfoot, whether by surgery or casting, produces adequate correction, the clubfoot deformity can recur as the child grows. In addition, sub-optimal initial correction can result in deviations in foot posture later in life. Assessment of foot kinematics during gait in a group

FIGURE 30.2 Clinical photograph of child with left clubfoot. OpenStax College (CC BY 3.0 ,https://creativecommons.org/licenses/by/3.0.) ,https://upload.wikimedia.org/wikipedia/commons/5/51/813_Clubfoot.jpg..

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of surgically treated children showed significantly increased forefoot extension and internal rotation, reduced range of sagittal plane motion of the hindfoot and increased inversion of the hindfoot compared to an age-matched control group [18]. Children treated with the Ponseti method have also been shown to demonstrate differences compared to a control population, including decreased range of motion at the ankle, internal foot progression angle, and reduced ankle power [19]. Under or over correction of a clubfoot deformity is likely to impact on the biomechanical function of the foot and lower limb as the child grows and may additionally result in pain during weight bearing activities. Currently, there are no clear guidelines on management of problematic clubfoot in the older child. Appropriate biomechanical assessment of the foot and ankle during functional tasks can aid the decision-making process, however, more research is needed in this area. Oblique and vertical talus is where the talus is plantarflexed compared to normal alignment in the sagittal plane and the navicular is dislocated dorsally. This in turn affects the position and orientation of the rest of the foot [20]. The hindfoot tends to be in equinus and valgus, with relative forefoot extension and external rotation and contraction of the dorsiflexor muscles. It commonly manifests as a rigid rocker-bottom foot [21]. Oblique or vertical talus can present as an isolated problem but may also be associated with neuromuscular conditions such as neurofibromatosis and arthrogryposis [22]. It is important to treat the condition early, to prevent progression of the deformity, particularly as the child learns to walk. There are distinctions between oblique and vertical talus; in patients with an oblique talus, full plantarflexion of the foot and ankle restores the normal talonavicular alignment, while in vertical talus this is not the case. Traditionally, both these conditions were treated surgically. However, more recently, these conditions, particularly oblique talus, have been treated successfully with serial casting and Achilles’ tenotomy [21]. Even if serial casting does not completely restore the anatomy and surgery is required, serial casting often makes the surgical re-alignment of the foot technically easier. Tarsal coalition describes a condition where two or more adjoining tarsal bones are abnormally connected by fibrous, cartilaginous, or bony tissue [21]. It is genetic in origin and is classified as a failure of segmentation [23]. It typically results in a rigid flat foot, with a planovalgus foot posture often associated with stiffness at the ankle and midfoot joints. The majority of coalitions involve the calcaneonavicular or talocalcaneal joints [21] (Figs. 30.3 and 30.4). Many coalitions are asymptomatic, and therefore do not necessarily require treatment. If symptoms develop, initial conservative treatment with orthoses is generally indicated. Non-responsiveness to conservative treatment may lead to surgical intervention to resect the coalition. There have been some studies assessing outcome following surgery using gait analysis. These have shown generally good clinical outcomes [23] but function can remain compromised [24] and the exact cause of this is still unknown. Accessory navicular occurs when the secondary ossification center of the navicular fails to fuse with the primary ossification center [5] (Fig. 30.5). It is mostly asymptomatic and often diagnosed incidentally when imaging the foot for other reasons. If prominent or associated with an idiopathic flat foot, it may cause pain due to pressure on the overlying skin. This is generally managed conservatively. Surgical intervention is rarely required [25] but if it is it often involves partially detaching some of the tibialis posterior tendon insertion footprint, resecting the accessory navicular and reefing the tendon footprint back onto the navicular. Few studies have investigated the effect of an accessory

FIGURE 30.3 Computed tomography (CT) 3D reconstruction image showing talonavicular coalition. Hellerhoff (CC BY-SA 3.0 ,https://creative,https://upload.wikimedia.org/wikipedia/commons/1/13/Kalkaneonavikulare_fibroese_Koalition_mit_Fraktur_ commons.org/licenses/by-sa/3.0.) Navikulare_-_CT_Volumen_Rendering_-_001_-_06.jpg..

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FIGURE 30.4 Plain radiograph (left) and CT image (right) of calcaneonavicular coalition. Tarsale Koalition 36jw—Roe OSG seitlich und CT Volumen rendering—001.jpg.

FIGURE 30.5 Plain radiograph showing accessory navicular of right foot. By Hellerhoff—Own work, CC BY-SA 3.0 ,https://commons.wikimedia. org/w/index.php?curid 5 39354099..

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navicular on gait or function. Kanatli et al. [26] used plantar pressure to assess the relationship between arch height index and presence and type of accessory navicular. They were unable to find any correlation and concluded that this condition is not associated with pes planus, however, the use of plantar pressure alone to assess arch height is questionable.

30.2.2 Developmental foot deformities Flexible flatfoot in children is a frequent presentation to health professionals. A flat medial longitudinal arch while standing is normal for the foot of children under the age of 6 years [10]. In most cases, the arch restores on walking or standing on tip toes due to the activation of, amongst others, the tibialis posterior muscle. Persistence of a flattened medial longitudinal arch beyond this age is less common and may be problematic. Most flexible flat feet in children are symptom free [27]. Concerns about development of future problems, or issues to do with shoe wear often trigger a visit to a health professional. Jack’s test (passive dorsiflexion of the hallux metatarsophalangeal joint) should restore the arch in the midfoot if the windlass mechanism is intact. Similarly, a heel rise test should restore the arch in idiopathic flat feet. These two tests are used to differentiate between flexible flat feet, in which the arch reconstitutes with either of these tests [28] or rigid flat feet, in which it does not. There are numerous studies in the literature assessing the effect of flexible flat feet in children on function and gait. These have found differences in foot and ankle kinematics [29] as well as in more proximal joints [13]. Differences between symptomatic and asymptomatic feet have also been noted [29]. The presence of flexible flat feet has also been shown to impact on quality of life [30]. While the foot remains flexible and symptom free, no treatment is indicated. In symptomatic feet, orthoses and stretching are often suggested. There is very little evidence in the literature however, indicating the efficacy of orthoses. In severe cases where conservative treatment has been ineffective, soft and or bony surgery may be indicated [28]. This is more often the case with structural rather than flexible flat feet. Rigid flat feet are discussed in more detail later. Juvenile bunion or hallux valgus is relatively uncommon in children [31]. It can arise due to increases in the intermetatarsal angle between the first and second metatarsals, or simply an increase in the metatarsophalangeal hallux valgus angle. In these patients, internal rotation of the first metatarsal is often also accompanied by relative flattening of the transverse forefoot arch and internal rotation of the hallux (Fig. 30.6). It can also be driven by laxity at the first tarsometarsal joint and this is typically the case in patients with generalized ligamentous laxity. It may also be present in neuromuscular conditions such as cerebral palsy [31]. Often the complaint is cosmetic but may also be painful in some cases. Asymptomatic cases do not require intervention. If the deformity becomes painful, Chell et al. [31] recommend that surgical intervention should be avoided prior to skeletal maturity, due to the high rate of recurrence. Conservative management usually involves alteration of shoe wear, orthoses aimed at restoring the transverse forefoot arch and longitudinal medial arch, and activity modification. Surgical correction depends on the driving deformity but typically involves corrective osteotomies of the first ray, with or without fusion of the first tarsometatarsal joint. While gait and function have been assessed in the adult population with hallux valgus, there is a scarcity of literature assessing children with this condition.

30.2.3 Foot pathologies associated with other conditions Rigid flat foot can be associated with other conditions such as cerebral palsy, some forms of muscular dystrophy, juvenile arthritis, as well as other congenital neurological and connective tissue disorders. In cases where this is asymptomatic and plantar areas that are usually non-weight bearing in neutral feet, such as medial midfoot, are adequately monitored, the flat foot itself may not require intervention. However, it is more common for rigid flat feet to be symptomatic, particularly when associated with a more systemic condition. A rigid flat foot disrupts locking and unlocking of the midfoot that is essential to the efficiency of walking gait. In normal feet, eversion of the subtalar joint unlocks the transverse tarsal joint and allows for a supple foot to accommodate the ground just after heel strike and inversion of the subtalar joint locks the transverse tarsal joint to allow for a stable hindfoot/midfoot for toe-off. In rigid flat feet, this delicate kinematic coupling is disrupted, and the arch cannot perform its normal roles of force transfer, accommodation of terrain, and shock attenuation. Kruger et al. [32] conducted a study on children with cerebral palsy and rigid flat foot. They found that children with cerebral palsy and flat foot walked significantly slower than typically developing peers. While there was a consistent finding of increased forefoot extension and external rotation in the cerebral palsy group, they found a range of deviations in the hindfoot coronal plane kinematics. They used radiographs to establish hindfoot axes and interestingly found eversion, neutral, and even

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FIGURE 30.6 Radiograph of foot showing prominent medial midfoot soft tissue shadow of a flat foot (single arrow) as well as prominent first metatarsal head of hallux valgus (double arrow). M Mifsud, Oxford University Hospitals NHS Trust.

inversion alignment during walking gait. While there is evidence in the literature that children with rigid flat foot walk differently to typically developing peers, there is less evidence to compare children within any given pathological group with and without a flat foot deformity. This makes it difficult to establish the unique effect of the flat foot deformity on walking gait. Symptomatic flat feet may be initially treated with orthoses and insoles. However, as the deformity becomes increasingly rigid or the adjacent joints start to be affected by the rigid flat foot, surgical intervention is often required. Pes Cavus or cavo-varus foot is a common presentation in several different pathologies. The most common pathology associated with this foot type is Charcot-Marie-Tooth (CMT) disease. CMT is a progressive peripheral neuropathy which causes muscle atrophy and impaired sensation [33]. The resulting muscle imbalances in the lower limbs with relative weakness of the foot evertors and ankle dorsiflexors leads to an equinocavovarus foot deformity which may be forefoot driven, due to flexion of the first metatarsal, or midfoot and hindfoot driven (Fig. 30.7). Clawed toes are often associated with this deformity. Early on, while the deformity remains flexible, the foot may be treated conservatively with orthoses and physiotherapy. Once the deformity becomes fixed, surgical intervention including soft and/or bony procedures may be indicated [34].

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FIGURE 30.7 Lateral radiograph of a right foot showing pes cavus. Note the prominent midfoot longitudinal arch and increased calcaneal pitch. Incidental finding of os peroneum too. Mikael Ha¨ggstro¨m File ,https://commons.wikimedia.org/wiki/File:Pes_cavus_and_os_peroneum_ on_lateral_foot_X-ray.jpg#/media/File:Pes_cavus_and_os_peroneum_on_lateral_foot_X-ray.jpg..

Cavo-varus foot deformities may also be present in children with cerebral palsy. In fact, it is the most common foot deformity in children with hemiplegic type cerebral palsy [35]. The spasticity of the ankle plantarflexors and foot invertors (which are relatively strong muscles compared to the foot evertors and ankle dorsiflexors) cause an imbalance of muscle forces around the ankle and foot and the foot adopts the typical position of hindfoot equinus and inversion, midfoot cavus, and forefoot varus and internal rotation. Gait analysis has been used to determine the functional impact of a cavo-varus foot deformity. In children and adults with CMT, increased plantarflexion at initial contact and push off, as well as changes more proximally at the hip and knee have been reported [36]. In children with cerebral palsy, a cavo-varus foot was also found to cause changes in hip and knee kinematics which resolved spontaneously when the foot was surgically corrected [37].

30.3

Functional assessment of the pediatric foot

The foot plays a vital role in enabling efficient gait and weight bearing function. Therefore, foot deformity in children can have a debilitating impact on activities of daily living and quality of life. One study examining surgically treated clubfoot in adults found the impact on quality of life to be equivalent, or even greater, than people with end-stage kidney disease or congestive heart failure [38]. Kothari et al. [30] demonstrated significantly lower scores on the Oxford Ankle Foot Questionnaire in children with flexible flat feet compared to those with neutral feet. To achieve optimal function, it is therefore important to be able to appropriately assess function of the foot and ankle during gait and other weight-bearing activities in order to guide intervention and assess outcomes. Functional assessment of the foot and ankle in children covers a range of different modalities including kinematics, kinetics, plantar pressure, and imaging. Combining different techniques, such as kinematics and plantar pressure, has also recently been proposed [39] . Kinematic assessment of the foot and ankle during gait has become more wide spread since the advent of multi-segment foot models. These models divide the foot into two or more segments to track motion during gait. A recent review found 65 papers that applied these models in a clinical context, however, only thirteen of these studies included children [40]. The smaller foot size and relative difficulty in palpating anatomical landmarks make the use of these models more challenging in children. Despite these difficulties, repeatability has been found to be comparable to that found when multi-segment models are applied to adult feet [41]. According to the review, five different multi-segment models have been applied to children (the Milwaukee, Oxford, Heidelberg, Rizzoli, and Shriners Hospital models), and have included populations with cerebral palsy, flat foot, clubfoot, forefoot varus, and juvenile idiopathic arthritis. The results were used to differentiate pathological populations from controls, classify foot types, assess relationship of foot deformity with quality of life, and determine the effect of different interventions [40]. It was concluded that multi-segment foot models may be useful in aiding clinical decision making, but longitudinal studies are required to verify this. Kinetic data has also been obtained from multi-segment foot models. However, this generally requires precise targeting of force plates during walking, such that the rear half of the foot is on one plate and the front half on the second plate. This was found to be viable and repeatable in a group of typically developing children [42] but is generally

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impractical for use in children with foot deformity and impaired gait. Combining a pressure and force plate could also provide this information [43] but is a cumbersome set up which most clinicians do not have access to. One potential solution could be to use musculoskeletal modeling with detailed foot segments to estimate forces and moments within the foot [44]. This has recently been suggested and applied to adult populations but has yet to be attempted with children. Plantar pressure analysis is another common method for biomechanical assessment of children with foot deformity. Two common pathologies where this assessment has been applied in children are clubfoot and cerebral palsy. In these populations, plantar pressure has been used to assess efficacy of specific interventions, to classify feet into different categories or levels of severity, to correlate plantar pressure findings with other measurement modalities or clinical assessments, and to characterize a specific group of patients by comparing the results to a control population [45]. The evidence suggests that plantar pressure measurements can be used to distinguish pathological from typically developing feet, and to quantify the outcome of interventions. However, there is currently very little evidence in the literature that links specific pressure parameters to relevant clinical parameters, or that predicts treatment outcomes in pediatric populations. This means that plantar pressure assessment can be useful to classify feet and quantify treatment outcomes, but its effectiveness in guiding treatment is less certain [45]. Further work is needed in this area to standardize protocols and reporting, and to clarify relationships between plantar pressure parameters (such as peak pressure values) with clinical impairments. Imaging techniques could also theoretically be used to assess foot function during activities such as gait. Fluoroscopy has been used to measure the motion of the foot bones during walking and running in adult populations but has yet to be applied in a pediatric population due to ethical considerations related to radiation exposure. Dynamic imaging with adequate spatial and temporal resolution but without the impact of radiation would be required to use this modality in children. Although still in its infancy, it is likely dynamic magnetic resonance imaging is the way forward in this regard. Static imaging data has been used in combination with dynamic measures such as multi-segment foot kinematics, to improve the accuracy of axis alignment [46]. Combining different modalities, such as kinematics with imaging or plantar pressure data, is a promising way forward to advance biomechanical and functional assessment of children’s feet. For example, tracking anatomical landmarks using a motion capture system has been used to improve the reliability and accuracy of masking plantar pressure images in typically developing children and children with clubfoot deformity [39]. It was found that using the combined approach allowed greater sensitivity in distinguishing clubfoot from the typically developing feet. In summary, there is evidence in the literature that multi-segment foot kinematics and plantar pressure data can be useful in assessing children’s feet, particularly in terms of classification of deformity and quantifying treatment outcomes. Some methods have been applied to adult feet, but these are not currently feasible for general application to children with foot deformity, for example dynamic fluoroscopic imaging and kinetic assessment. Combining multiple techniques is likely to provide further insight into the biomechanical function of both typically developing and pathological feet.

30.4

Areas for future research

Most research in the field of biomechanical assessment of children’s feet is composed of cross-sectional studies. The majority of these compare pathological to typically developing cohorts. There are very few studies that follow children over the longer term as they grow and develop. There is a need for long-term studies that document the changes in children’s feet; in particular how they respond to extrinsic and intrinsic influences. Studies that document the effect of interventions such as surgery and casting on the biomechanics and function of children’s feet, as well as the subsequent effects on more proximal joints, are needed. It would also be beneficial to know the long-term effects of restriction of motion of the foot, for example, through orthotics or shoe-wear, on the growth, development and function of children’s feet. Developing innovative methods for biomechanical assessment of children’s feet would also be beneficial in improving our understanding of function and therefore promoting optimal management. This could involve integrating different techniques (such as imaging and motion analysis) or applying techniques in a different context. Advanced modeling techniques (including musculoskeletal modeling) could provide an avenue for including assessment of forces and moments within the foot. In conclusion, it is important to understand the pediatric foot to promote optimal growth and appropriately manage foot-related pathology. Pediatric feet differ from adult feet in both structure and function, and this needs to be assessed and treated accordingly. While much progress has been made in understanding the biomechanics of children’s feet,

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there are still areas in need of research. Improving techniques for measuring children’s feet could help to provide better treatment outcomes and therefore improve the quality of life for children with foot pathology.

References [1] Currey JD, Butler G. The mechanical properties of bone tissue in children. J Bone Jt Surg Ser A 1975;57-A:810 14. Available from: https:// doi.org/10.2106/00004623-197557060-00015. [2] Gould N, Moreland M, Trevino S, Alvarez R, Fenwick J, Bach N. Foot growth in children age one to five years. Foot Ankle Int 1990;10:211 13. Available from: https://doi.org/10.1177/107110079001000404. [3] Dimeglio A, Stanitski CL. Growth in pediatric orthopaedics. J Pediatr Orthop 2001;21:549 55. Available from: https://doi.org/10.1097/ 00004694-200107000-00026. [4] Drennan JC. Anatomy. In: McCarthy J, Drennan J, editors. Child’s foot ankle. 2nd ed Philadelphia: Wolters Kluwer & Lippincott Williams & Wilkins; 2010. [5] Sarrafin S, Kelikian A. Development of the foot and ankle. In: Kelikian A, editor. Anatomy of the foot and ankle. Philadelphia: Wolters Kluwer & Lippincott Williams & Wilkins; 2011. [6] Lisowski FP. Angular growth changes and comparisons in the primate talus. Folia Primatol 1974;7:81 97. Available from: https://doi.org/ 10.1159/000155111. [7] Uden H, Scharfbillig R, Causby R. The typically developing paediatric foot: how flat should it be? A systematic review. J Foot Ankle Res 2017;. Available from: https://doi.org/10.1186/s13047-017-0218-1. [8] Bernhardt DB. Prenatal and postnatal growth and development of the foot and ankle. Phys Ther 1988;68:1831 9. Available from: https://doi. org/10.1093/ptj/68.12.1831. [9] Cappello T, Song KM. Determining treatment of flatfeet in children. Curr Opin Pediatr 1998;18:142 9. Available from: https://doi.org/ 10.1007/978-3-319-33622-0_22. [10] Mickle KJ, Steele JR, Munro BJ. The feet of overweight and obese young children: are they flat or fat? Obesity 2006;14:1949 53. Available from: https://doi.org/10.1038/oby.2006.227. [11] Fritsch H. Sectional anatomy of connective tissue structures in the hindfoot of the newborn child and the adult. Anat Rec 1996;246:147 54. Available from: https://doi.org/10.1002/(SICI)1097-0185(199609)246:1%3E147::AID-AR16%3C3.0.CO;2-P. [12] Hennig EM, Staats A, Rosenbaum D. Plantar pressure distribution patterns of young school children in comparison to adults. Foot Ankle Int 1994;15:35 40. Available from: https://doi.org/10.1177/107110079401500107. [13] Kothari A, Dixon PC, Stebbins J, Zavatsky AB, Theologis T. Are flexible flat feet associated with proximal joint problems in children? Gait Posture 2016;45:204 10. Available from: https://doi.org/10.1016/j.gaitpost.2016.02.008. [14] O’Shea RM, Sabatini CS. What is new in idiopathic clubfoot? Curr Rev Musculoskelet Med 2016;9:470 7. Available from: https://doi.org/ 10.1007/s12178-016-9375-2. [15] Engell V, Nielsen J, Damborg F, Kyvik KO, Thomsen K, Pedersen NW, et al. Heritability of clubfoot: a twin study. J Child Orthop 2014;8:37 41. Available from: https://doi.org/10.1007/s11832-014-0562-7. [16] Faizan M, Jilani LZ, Abbas M, Zahid M, Asif N. Management of idiopathic clubfoot by Ponseti technique in children presenting after one year of age. J Foot Ankle Surg 2015;54:967 72. Available from: https://doi.org/10.1053/j.jfas.2014.05.009. [17] Zhao D, Li H, Zhao L, Liu J, Wu Z, Jin F. Results of clubfoot management using the Ponseti method: do the details matter? A systematic review. Clin Orthop Relat Res 2014;472:1329 36. Available from: https://doi.org/10.1007/s11999-014-3463-7. [18] Theologis TN, Harrington ME, Thompson N, Benson MKD. Dynamic foot movement in children treated for congenital talipes equinovarus. J Bone Jt Surg 2003;85:572 7. Available from: https://doi.org/10.1302/0301-620X.85B4.13696. [19] Mindler GT, Kranzl A, Lipkowski CAM, Ganger R, Radler C. Results of gait analysis including the oxford foot model in children with clubfoot treated with the ponseti method. J Bone Jt Surg Am 2014;96:1593 9. Available from: https://doi.org/10.2106/JBJS.M.01603. [20] Coleman S, Stelling F, Jarrett J. Pathomechanics and treatment of congenital vertical talus. Clin Orthop Relat Res 1970;70:62 72. [21] Dare DM, Dodwell ER. Pediatric flatfoot: cause, epidemiology, assessment, and treatment. Curr Opin Pediatr 2014;26:93 100. Available from: https://doi.org/10.1097/MOP.0000000000000039. [22] Drennan JC. Congenital vertical talus. J Bone Jt Surg 1995;77:1916 23. [23] Lemley F, Berlet G, Hill K, Philbin T, Isaac B, Lee T. Current concepts review: tarsal coalition. Foot Ankle Int 2006;27:1163 9. Available from: https://doi.org/10.1177/107110070602701229. [24] Kitaoka HB, Wikenheiser MA, Shaughnessy WJ, An KN. Gait abnormalities following resection of talocalcaneal coalition. J Bone Jt Surg [Am] 1997;79:369 74. [25] Knapik DM, Guraya SS, Conry KT, Cooperman DR, Liu RW. Longitudinal radiographic behavior of accessory navicular in pediatric patients. J Child Orthop 2016;10:685 9. Available from: https://doi.org/10.1007/s11832-016-0777-x. [26] Kanatli U, Yetkin H, Yalcin N. The relationship between accessory navicular and medial longitudinal arch: evaluation with a plantar pressure distribution measurement system. Foot Ankle Int 2003;24:486 9. Available from: https://doi.org/10.1177/107110070302400606. [27] Chytas A, Morakis E. Foot disorders in children. Surg (U Kingd) 2016;35:48 51. Available from: https://doi.org/10.1016/j.mpsur.2016.10.007. [28] Mosca VS. Flexible flatfoot in children and adolescents. J Child Orthop 2010;4:107 21. Available from: https://doi.org/10.1007/s11832-0100239-9.

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[29] Kerr CM, Zavatsky AB, Theologis T, Stebbins J. Kinematic differences between neutral and flat feet with and without symptoms as measured by the Oxford foot model. Gait Posture 2019;67:213 18. Available from: https://doi.org/10.1016/j.gaitpost.2018.10.015. [30] Kothari A, Stebbins J, Zavatsky B, Theologis T. Health-related quality of life in children with flexible flatfeet: a cross-sectional study. J Child Orthop 2014;8:489 96. Available from: https://doi.org/10.1007/s11832-014-0621-0. [31] Chell J, Dhar S. Pediatric hallux valgus. Foot Ankle Clin 2014;19:235 43. Available from: https://doi.org/10.1016/j.fcl.2014.02.007. [32] Kruger K, Konop K, Krzak J, Graf A, Altiok H, Smith P, et al. Segmental kinematic analysis of planovalgus feet during walking in children with cerebral palsy. Gait Posture 2017;54:277 83. Available from: https://doi.org/10.1016/j.gaitpost.2017.03.020. [33] Brewerton D, Sandifer P, Sweetnam D. Idiopathic pes cavus: an investigation into its aetiology. Br Med J 1963;2:659 61. [34] Maranho DA, Volpon JB. Acquired pes cavus in Charcot-Marie-Tooth Disease. Rev Bras Ortop Engl (Ed.) 2009;44:479 86. Available from: https://doi.org/10.1016/S2255-4971(15)30144-0. [35] Wren TAL, Rethlefsen S, Kay RM. Prevalence of specific gait abnormalities in children with cerebral palsy: influence of cerebral palsy subtype, age, and previous surgery. J Pediatr Orthop 2005;25:79 83. Available from: https://doi.org/10.1097/00004694-200501000-00018. [36] Newman CJ, Walsh M, O’Sullivan R, Jenkinson A, Bennett D, Lynch B, et al. The characteristics of gait in Charcot-Marie-Tooth disease types I and II. Gait Posture 2007;26:120 7. Available from: https://doi.org/10.1016/j.gaitpost.2006.08.006. [37] Stebbins J, Harrington M, Thompson N, Zavatsky A, Theologis T. Gait compensations caused by foot deformity in cerebral palsy. Gait Posture 2010;32. Available from: https://doi.org/10.1016/j.gaitpost.2010.05.006. [38] Dobbs MB, Nunley R, Schoenecker PL. Long-term follow-up of patients with clubfeet treated with extensive soft-tissue release. J Bone Jt Surg Ser A 2006;88:986 96. Available from: https://doi.org/10.2106/JBJS.E.00114. [39] Giacomozzi C, Stebbins JA. Anatomical masking of pressure footprints based on the Oxford Foot Model: validation and clinical relevance. Gait Posture 2017;53:131 8. Available from: https://doi.org/10.1016/j.gaitpost.2016.12.022. [40] Leardini A, Stebbins JA, Caravaggi P, Theologis T. Multi-segment foot models and their use in clinical populations. Gait Posture 2019; in press. [41] McCahill J, Stebbins J, Koning B, Harlaar J, Theologis T. Repeatability of the Oxford Foot Model in children with foot deformity. Gait Posture 2018;61:86 9. Available from: https://doi.org/10.1016/j.gaitpost.2017.12.023. [42] Bruening DA, Cooney KM, Buczek FL. Analysis of a kinetic multi-segment foot model. Part I: model repeatability and kinematic validity. Gait Posture 2012;35:529 34. Available from: https://doi.org/10.1016/j.gaitpost.2011.10.363. [43] Giacomozzi C, Macellari V. Piezo-dynamometric platform for a more complete analysis of foot-to- floor interaction. IEEE Trans Rehabil Eng 1997;5:322 30. Available from: https://doi.org/10.1109/86.650285. [44] Fluit R, Andersen MS, Kolk S, Verdonschot N, Koopman HFJM. Prediction of ground reaction forces and moments during various activities of daily living. J Biomech 2014;47:2321 9. Available from: https://doi.org/10.1016/j.jbiomech.2014.04.030. [45] Stebbins J. Assessing clubfoot and cerebral palsy by pedobarography. In: Muller B, Wolf SI, editors. Handbook of human motion. Cham: Springer; 2018. p. 727 39. Available from: http://doi.org/10.1007/978-3-319-14418-4_37. [46] Kidder SM, Abuzzahab FSJ, Harris GF, Johnson JE. Clinical validation of a system for the analysis of pediatric foot and ankle kinematics during gait. IEEE Trans Rehabil Eng 1996;4:25 32. Available from: https://doi.org/10.1109/IEMBS.1998.744918.

Chapter 31

Neurological Foot Pathology Morgan E. Leslie1 and Joseph M. Iaquinto1,2,3 1

RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 2Department of

Mechanical Engineering, University of Washington, Seattle, WA, United States, 3Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States

Abstract This chapter provides an overview of pathologies involving neurological mechanisms affecting foot and ankle biomechanics including stroke, cerebral palsy, toe walking, peripheral neuropathy, foot drop, tarsal tunnel syndrome, Morton’s neuroma, Charcot foot, Charcot-Marie-Tooth disease, Friedreich’s ataxia, and poliomyelitis. A brief literature summary of their etiology, impact on biomechanics, and treatment is included. Each section discusses the variations in biomechanics, gait, and neuromuscular anatomy that might occur with each pathology as it relates to the foot and ankle. The implications on quality of life and potential rehabilitative outcomes are reported. A variety of treatments (standard of care, in research, novel developments) for each pathology are described along with brief discussions on the efficacies of each treatment.

31.1

Introduction

The appropriate functioning of the neuromuscular control system is required for effective performance of the foot and ankle. In the case of diseases affecting the nervous system, the foot and ankle can often be altered. This chapter provides a brief overview of several common neurological pathologies affecting the foot and ankle. Neurological pathologies in the foot and ankle can have extensive impacts on the rest of the body and its ability to function effectively, thereby impacting quality of life. Individuals with issues originating in the foot and ankle tend to have multiple comorbidities and impairments in alignment, weight distribution, motion, and neuromuscular performance during static and/or dynamic activities [1]. 30% 37% of adults older than 45 years old are affected by foot pain [1]. Despite this, there are relatively few studies evaluating foot-specific conditions and possible interventions; however, with improving technology in data collection methods (i.e. motion capture, medical imaging, and plantar pressure), significant enhancements have begun to be made in the last decade in the analysis of foot and ankle conditions, but studies on clinical interventions are still lacking [1].

31.2

Stroke

31.2.1 Pathology related to the musculoskeletal system Stroke can affect various functions of the body including those governed by the musculoskeletal system, causing impairments and resulting in compensatory mechanisms. Musculoskeletal dysfunction after stroke can be explained by many factors, including muscle atrophy and weakness due to compensatory mechanisms as well as reduced activity. Normally, there are five distinct modules of synergistic muscular control involved in gait; however, stroke subjects with more gait asymmetry have fewer distinct modules as some can be merged as a result of compensation by the central nervous system in response to muscle weakness and decreased muscle control [2]. These complications can alter the physiology of the body, including that of the foot and ankle, which can have a negative impact on mobility. Deformities resulting from stroke in the foot and ankle and its function crucial to weightbearing and balance can have negative implications on gait. When compared to healthy controls, stroke patients exhibit deficits in muscle size Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00047-0 © 2023 Elsevier Inc. All rights reserved.

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and strength in both limbs [3]. Hemiplegia, paralysis on one side of the body, and hemiparesis, one-sided weakness, are common disabilities caused by stroke [4]. While there exists a great deal of research and literature focusing on hemiplegic gait after stroke, most of this focuses on hip and knee motion; however, deficits in the foot and ankle may have an equal or greater impact on the quality of gait. Post-stroke survivors experience approximately 52% and 36% reduced strength in their plantarflexors and knee extensor muscles respectively on the affected side when compared to the unaffected [3]. Foot deformities that do affect walking quality is reported in approximately 50% of chronic stroke survivors. About 30% of stroke patients are said to experience abnormal and asymmetric foot posture while standing [5].

31.2.2 Impact on kinematics Abnormal ankle joint movement—due to its important role in absorbing impact, body advancement, and balance—contributes significantly to gait disorders. The most significant abnormality is reduced hindfoot plantarflexion motion and duration after initial contact and late stance [5]. Stroke survivors with difficulty walking have altered neural control and modular organization. Reduced plantarflexor function demonstrates an impact on the body’s ability to execute proper forefoot and heel rocking motions for forward propulsion of weightbearing. Approximately 67% of variance in gait speeds can be attributed to plantarflexor function impairment [3]. Hindfoot movements are most strongly correlated to walking ability; therefore, stroke survivors who experience reduced plantarflexion, inversion, or adduction in their hindfoot tend to be considered “household walkers” and are thus limited in their walking ability [5]. Discrepancies in foot motion, organized by foot segment, planes of motion and distinct phases of motion, that have been observed are as follows [5] (Table 31.1). A graphical representation of the angles of planar motion in different foot segments in stroke patients as compared to healthy controls is presented (Fig. 31.1).

TABLE 31.1 The kinematic differences between stroke survivors compared to healthy controls. Segment

Plane of motion

Kinematic difference

Whole foota

Sagittal

First phase: 1.8 degrees less plantarflexion, shorter duration Second phase: comparable range of motion (ROM) and maximum dorsiflexion, longer duration Last phase: reduced ROM and maximum plantarflexion

Frontal

No significant differences in inv/eversion across the three phases

Transverse

4 degrees reduced adduction, followed by maximum abduction (normal pattern involves abduction then adduction)

Sagittal

First phase: 2.1 degrees reduction in plantarflexion Last phase: less plantarflexion ROM, reduced maximum plantarflexion at toe off

Frontal

3 degrees less ROM, greater maximum eversion Last phase: Less inversion at toe off

Transverse

2.6 degrees less ROM First phase: more adduction at initial contact, less abduction and delayed timing of abduction Second phase: Overall less movement during adduction

Mid-foot

All

No significant differences across all planes of motion

Forefoot

Sagittal

First and Second phases: no significant differences Last phase: 2.7 degrees less ROM and delayed maximum dorsiflexion

Frontal

No significant differences

Transverse

More abduction over entire stance phase Last phase: less ROM and maximum adduction in late stance

Hindfoot

Analysis based on a multi-segment foot model. Each plane has distinct phases of motion—sagittal: three phases of dorsi/plantarflexion; frontal: two phases of inv/eversion; transverse: two phases of add/abduction [5]. a Refs. [2,3,6] discuss similar trends of gait impairment, but they all focus only on the whole foot.

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Sagial Plane Movements

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FIGURE 31.1 Planar (sagittal and frontal) movements of foot and ankle (segments A D) during stance phase in stroke (mean 5 dash line; 1 / 2 1 SD 5 gray band) versus control (mean 5 solid line; 1 / 2 1 SD 5 error bar) [5].

31.2.3 Impact on foot function Structural deformity and loss of function of the foot can have compounding impacts on the ability of the foot to maintain balance, causing alterations in gait and resulting in musculoskeletal pain. People living with post-stroke effects experience decreased sensation of the foot, particularly at the plantar aspect of the first metatarsophalangeal joint (MPJ) [6]. The body’s ability to maintain foot posture stability and sense of foot position is understood to be largely due to sensory input of the plantar aspect of the foot. The first MPJ of stroke survivors’ paretic side also exhibits reduced range of motion (ROM) when compared to their unaffected side. A normal gait cycle requires 65 degree MPJ ROM, which is not achieved by sample stroke populations (at an average of 27.0 degree on the affected side) [6]. Dorsiflexion is slightly lower (9.5 degree in stroke subjects vs 10.5 degree in controls) which could be caused by stiff plantarflexors and weak dorsiflexors [6]. Swing phase requires at least 5 10 degree of dorsiflexion to achieve foot clearance, and

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FIGURE 31.2 Examples of stroke-caused foot and ankle impairments including: (A) plantarflexed, inverted foot from muscular contractions; (B) contraction of the toes upon weight bearing; and (C) pressure sores from AFOs used for treatment [7].

more than 50% of sample stroke survivors exhibit dorsiflexion below this range during swing phase on the paretic side [6]. Dorsiflexion is important in foot clearance during gait, but it can also be associated with postural stability and foot and ankle function. Stroke can cause various foot and ankle muscular contractures (Fig. 31.2).

31.2.4 Clinical treatment Stroke survivors are often provided with an ankle-foot orthosis (AFO); however, this device only provides support to deficits in motion in the sagittal plane. An understanding of the deficits seen in the distinct aspects of the foot might provide insight on improvement in clinical treatment. Instead of a traditional AFO that focuses on improving dorsiflexion, multi-segment analysis would suggest the use of a rollover orthotic to help improve forward progression due to defective heel and forefoot rocker functions from impaired hindfoot plantarflexion [5]. Common clinical belief suggests that equinovarus is the most frequently seen foot deformity post-stroke; however, multi-segment analysis shows excessive hindfoot eversion is also prevalent, which would suggest the need for support to the medial side of the foot (by way of arch supports or heel wedges) [5]. Rehabilitation is a critical part of motor recovery after stroke. Many rehabilitation tasks consist of the survivors relearning motor activities (i.e., sitting, standing up, standing, and walking). These tasks consist of the patient receiving biological feedback from sensory systems. Traditionally, therapists provided feedback in the form of a “coach” role, but recent advances in technology have provided the ability to deliver biofeedback to the patient concurrently to the task being done [8]. A systemic review by Stanton et al. that investigated the effects of biofeedback on stroke rehabilitation of the lower limb provides evidence that biofeedback produces results superior to those from traditional therapy and/or therapist verbal feedback. The most commonly used type of biofeedback in gait rehabilitation has been electromyographic, kinematic, and measures from robot-assisted gait training [9]. Another effective intervention to improve muscle strength, balance, gait, and motor function in the lower limb in hemiplegic stroke patients is mirror therapy with afferent electrical stimulation [10].

31.3

Cerebral palsy

31.3.1 Definition Cerebral palsy (CP) occurs in 2 3 out of 1000 live births, making it the most common physical pediatric disability [11]. CP is a group of non-progressive and permanent disorders that affects movement, posture, and balance, causing limitations in activity, caused by an injury to an infantile or fetal brain, disrupting normal development [11,12]. In an International Workshop on Definition and Classification of CP in 2004, due to various factors (including advancements in brain imaging technology, improved understanding of brain development, and recognition that motor deficits should

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not be an exclusive focus), previous definitions of CP were considered inadequate, and the term should be considered a description of clinical characteristics [12]. While there are some risk factors, most occurring during the perinatal period, 80% of CP cases are considered idiopathic [11]. CP as a diagnosis is classified by the movement disorders such as spasticity (the most common, seen in 80% of cases), ataxia (poor coordination), dyskinesia (loss of motor control), or a combination of multiple [11]. The motor disorders characteristic to CP are commonly accompanied by other developmental disorders, such as impairments in sensation, cognition (50% of cases), perception, communication (25%), and behavior (20% 25%) as well as epilepsy (25%) and other musculoskeletal deficits, including hip dysplasia (33%) [11,12]. The musculoskeletal pathology of CP includes the inability for skeletal muscle to grow longitudinally normally [13]. Muscle growth requires physiological loading and stretching of relaxed muscle, but due to the spasticity in children with CP, their muscle cannot relax during activity and activity is difficult because of weakness and imbalance; therefore, activity levels are reduced [13].

31.3.2 Structural deformities and gait deviations Up to 93% of pediatric CP patients have some kind of foot and ankle deformity [14], with equinus, equinoplanovalgus, and equinocavovarus being the three most common [15]. Pedobarographic loading patterns are distinctive for (A) equinoplanovalgus and (B) equinocavovarus feet (Fig. 31.3). Common foot and ankle deformities seen in CP cases can cause various disruptions along the gait cycle such as: non-existent heel strike at initial contact, compromised rocker and shock absorption functions in loading response, lack of dorsiflexion in swing phase and clearance in midswing, and incorrect foot and ankle positioning in terminal swing [15]. The deformities and their effect on gait can vary depending on the extent of the effects on the body. The type of brain lesion and its location produces different effects: a one-sided and central lesion in the motor cortex can cause hemiplegic (unilateral) CP, a lesion in the periventricular region of the brain affects both lower limbs (diplegic CP), and an injury to the brainstem due to widespread oxygen-deprivation can lead to quadriplegic (total body) CP [14]. Those with hemiplegic CP can walk independently, those with diplegic CP often need assistive devices to walk, and rarely are those with quadriplegic CP ambulatory [13]. Across the various types of CP, the most common gait abnormalities are stiff knee during swing phase (80%), crouch gait (69%), increased hip flexion (65%), intoeing (64%), and equinus (61%) [16,17]. Upon investigating each type of CP, there are differences in the incidences of certain gait deviations and foot abnormalities across groups and region of paralysis [16,17].

(A)

(B)

FF

FF

MF

MF

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HF

FIGURE 31.3 Pedobarographic data of two CP patients with (A) equinoplanovalgus deformity and (B) equinocavovarus deformity. The horizontal black lines divide the foot into three segments: hindfoot (HF), midfoot (MF), and forefoot (FF). The red line denotes the mean value for normal center of pressure (COP), while the black vertical black lines represent 1 / 2 1 and 2 SD from the mean COP [15].

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31.3.3 Treatment Management of deformities in the foot and ankle in children with CP is dictated by the severity of the deformity, which can be classified in levels: level I including deformities caused by spasticity and imbalance of muscles, level II deformities are caused by fixed soft tissue imbalance with no skeletal abnormality, and level III consisting of skeletal malalignment with fixed soft tissue imbalance [15]. Level I deformities are treated by pharmaceutical (the most common being phenol and botulinum toxin A (BTX-A)) or neurosurgical interventions [13 15]. Phenol is only used to improve adductor and elbow flexor spasticity, while BTX-A has broader affects which include management of the spasticity of the adductors (causing scissoring), hamstrings (causing crouching gait), plantarflexors and dorsiflexors (causing ankle equinus), and the upper limbs. In flexible, mild deformities, orthoses can be used to help improve gait efficacy. AFOs are common for management of sagittal plane abnormalities (like ankle equinus) by improving foot clearance during swing phase [14]. As deformities progress and become fixed or rigid (level II and level III), orthopedic surgeries are recommended, specifically muscle tendon lengthening and a combination of soft tissue and skeletal procedures [15]. Orthopedic surgery can be done to correct deformities to improve function, and it appears to be effective in decreasing the chances of acquiring equinus, intoeing, and ankle varus, but it seems to increase the chances of developing crouched gait, stiff knee gait, calcaneus gait, and out-toeing [16,17]. 74% of a group of hemiplegic children who had previous surgery had crouch gait, while only 37% developed it without surgical history [16,17]. There is no cure for CP, thus the main goal of these treatment options are to preserve energy and aim towards painless gait in ambulatory children, and to avoid skin irritation and pain in non-ambulatory children [14].

31.4

Toe walking

31.4.1 Diagnosis and etiology Toe walking is characterized by a lack of heel strike and walking on the toes or forefoot [18]. Toe walking is considered to be a part of the natural development of gait [18,19]. A typical heel-toe pattern should develop by 24 42 months [20]. However, persistent toe walking beyond the developmental stage can be a sign of an underlying structural abnormality (i.e., contracted tendon, limb length discrepancy, ankylosing spondylisis) or neuromuscular disorders (i.e., autism, Charcot-Marie-Tooth [CMT], CP, Duchenne muscular dystrophy) [18,19]. The highest number of toe walkers is found in children with autism spectrum disorder, and the second highest being found in children with communication disorders, though no concrete physiological cause is known. Persistent toe walking is found in these cases due to sensory processing disorders [20]. Duchenne muscular dystrophy (DMD) and CP are the most common progressive and non-progressive pediatric neuromuscular disorders, respectively, and toe walking is seen as a characteristic in both [20]. In CP, toe walking is generally attributed to tight and spastic plantarflexors and in these cases children are unable to achieve flat foot stance [20]. For children with DMD, toe walking is a compensatory strategy that allows them to ambulate despite lower limb weakness and knee instability [20]. Idiopathic toe walking (ITW) is diagnosed upon exclusion of all other diagnoses after thorough physical examinations and neurological/psychiatric diagnostic tests (summarized by the “Toe Walking Tool,” a 28-question evaluation). Patients can usually walk on flat feet without difficulty when asked [19,20]. Tendon contractures which could lead to limited ankle ROM, which might be a cause of ITW or could be a result of persistent toe walking, but this relationship is uncertain [19,20].

31.4.2 Biomechanical and musculoskeletal function Children with ITW have abnormal function and structure of the triceps surae muscle, characterized by myopathic and atrophic fibers, smaller and more abundant quantities of type I fibers when compared to type II, and premature activation during gait [21]. There is a 25% reduction in tibialis anterior muscle strength and three times more restricted passive ankle dorsiflexion in children with ITW when compared to healthy controls [21]. These abnormalities in muscle strength and ROM can contribute to poor performance in balance and development of the equinus gait pattern characteristic of ITW [21]. Normal ankle kinematics during gait includes the presence of three rocker phases: ankle plantarflexion lowers the foot to be flat on the ground at heel strike via contraction of the tibialis anterior; dorsiflexion through contraction of the gastrocnemius-soleus complex to progress the body forward over the foot; and ankle plantarflexion during push-off at the end of stance phase powered by contraction of the gastrocnemius and soleus [19]. During swing phase, the ankle is dorsiflexed by the tibialis anterior muscle to permit foot clearance [19]. In subjects with ITW, studies have shown the following changes in all three rockers: absent first rocker, decreased dorsiflexion during second, and increased plantarflexion during third rocker and swing phase [19]. When cued, ITW subjects can walk with a

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FIGURE 31.4 Kinetic patterns of Idiopathic toe walking (ITW) based on severity (type 1: mild (solid black line), type 2: moderate (red dashed line), type 3: severe (green dashed line) [22].

semi-normal gait pattern, but most cannot achieve complete normalcy [19]. Kinetic patterns (mild, moderate, and severe) during the gait cycle are observed in subjects with ITW (Fig. 31.4).

31.4.3 Treatment When toe walking is a result of a neuromuscular disorder, recommended treatments of that disorder must be followed, and toe walking is managed accordingly. For ITW, the decision to get treatment varies since there is little evidence supporting the correlation between poor functional outcomes after toe walking into adolescence, but treatment can be justified due to the psychological effect of ITW on parents and children [19]. Physical therapy or braces/splints to stretch the heel cord and triceps surae complex or even pharmaceutical intervention (botulinum toxin injections) can help improve ankle ROM, but the results of these nonsurgical interventions are difficult to evaluate due to the tendency to use multiple treatments in tandem [19]. Various studies have investigated the amount of subjects that showed improvements in gait, with varying results: 79% had a normal gait after casting/nighttime bracing and passive stretching regimens [23], 50% after observation or 6 weeks of casting, though only 22% returned to normal [24], and 66% had improvements after casting only [25]. Surgical treatment of ITW can be performed to help lengthen parts or all of the triceps surae muscle-tendon complex when dorsiflexion is very limited (fixed ankle equinus contractures) or when toe walking continues to recur even after nonsurgical treatments [19]. Surgical correction tends to have good outcomes: Eastwood et all saw 72% improved or normal gait after surgery compared to 51% after only casting and observation [24] and Stricker and Angulo saw 67% parental satisfaction after surgery versus 25% after only casting and observation [26]. Optimal treatment for ITW is unclear, but many conclude that decision on treatment is dependent on age (ITW might be outgrown overtime, so treatment should only be considered in older children) and prevalence and severity of tendon/muscular contractures [18,19].

31.5

Peripheral neuropathy

31.5.1 Background Proper gait requires a fully functional musculoskeletal system, sufficient muscular strength, and sensory input [27]. Loss of sensory input in the feet can modify the ability of the body to adjust motor patterns and muscle response to accomplish a task effectively [27]. Foot deformities can be caused by peripheral neuropathy (Fig. 31.5). Lack of sensory input in the feet can be caused by the onset of peripheral neuropathy, which is an encompassing term for disorders that affect the peripheral nervous system in various patterns [29]. The most common pattern is called distal sensory polyneuropathy (DSP), also known as diabetic peripheral neuropathy (DPN), which affects more than 50% of people with diabetes [29,30]. DSP consists of an injury to a peripheral nerve resulting in predominantly distally located sensor loss, pain, and weakness which can cause gait instability and increase the risk of falling, foot ulceration,

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Peripheral neuropathy

Peripheral vascular disease

Charcot foot

Glove and stocking neuropathy

Neuropathic ulcer

Dry, cracked skin Callus

Proximal arterial occlusion

Loss of leg hair

Absent pulses Cold feet

Clawing of toes

Digital gangrene

FIGURE 31.5 Charcot foot, ulcers, and clawed toes can be the result of peripheral neuropathy due to the disruption of normal muscle function and biomechanics. A Charcot foot occurs in the presence of neuropathy and results in bony destruction. Clawing of the toes can be caused by imbalance in muscle function and presence of muscle atrophy, leading to increased pressure on the metatarsal heads. These deviations in normal biomechanics can lead to ulcerations [28].

and amputations [30]. The changes in biomechanics caused by DSP can cause areas of elevated plantar pressures on the foot, in particular at the first MPJ, which can contribute to the development of ulcerations [30].

31.5.2 Musculoskeletal and movement implications When compared to healthy controls, people with DSP tend to walk slower (0.92 vs 1.14 m/s), with a wider support base, shorter steps (1.19 vs 1.34 m), limited knee and ankle mobility, lower plantar flexion moment and power, and more time spent in stance phase (61.07% vs 59.4% of the gait cycle) [27,30]. People with DPN displayed delays in the activation of their tibialis anterior and vastus lateralis muscles, resulting in decreased ankle efficiency [27]. DSP subjects showed only 55% of the mean plantarflexor torque, 59% of the ankle strength, and 83% of the knee strength found in healthy demographic-matched controls [27]. Muscle atrophy is also evident in people with DSP, with a 32% reduction of muscle volume when compared to controls [27]. DSP patients have a thicker plantar aponeurosis (4.2 mm vs 3.6 mm) and flexor hallucis longus (4.8 mm vs 4.3 mm) [27]. DPN tends to lead to less ankle motion (17.9 degrees vs 28.4 degrees 31 degrees), first MTP joint mobility (35.3 degrees vs 59.4 degrees 62 degrees), and subtalar joint mobility (18 degrees vs 35 degrees) [27]. Increased mean plantar pressures were seen in subjects with DSP in the hindfoot (39.84 vs 32.98 N/cm2), midfoot (24.6 vs 16.1 N/cm2), first MTP joint (46.5 vs 27.3 N/cm2), forefoot (47.17 vs 37.83 N/cm2), hallux (39.1 vs 33.05 N/cm2) [30].

31.6

Foot drop

31.6.1 Pathology Foot drop is a condition resulting from the disruption of the neural pathway to the dorsiflexor muscles of the foot. There are many potential etiologies, including trauma, disease or other causes, and depending on the process, foot drop can be a uni- or bilateral condition. This condition presents as weakness of the dorsiflexors such that the individual

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cannot lift their foot against gravity [31]. Those with this condition may not be able to support their own weight [32]. Additionally, depending on cause, foot drop can be continuous or spastic [33].

31.6.2 Impact on biomechanics Compensation for foot drop tends to be focused on preventing the toes from catching or dragging along the ground. Compensation is accomplished with hyperflexion of the hip and knee to raise the foot higher off the ground and clear the dangling foot from any obstacle. This condition and its compensation can lead to various system muscle and gait imbalances, and a shortening of the Achilles tendon [31]. A comparison of the biomechanics of the foot and ankle during the gait cycle between a normal foot and one with foot drop demonstrates the pathological changes (Fig. 31.6).

31.6.3 Treatment A wide range of treatment options are possible, depending on the cause and particulars of this condition. Treatments range from AFOs and other braces, including powered assistance; electrical stimulation devices; neurological surgery/ repair; and tendon transfer. Passive AFO devices mitigate the impact of foot drop by providing positional assistance to the ankle during gait, removing the “drop” component of foot drop. This reduces the compensation needed and thus the negative impacts of long-term systemic compensation but does not address the issue directly. More complex bracing works to restore the function of foot position during gait, through either unpowered or powered mechanical assistance with dorsiflexion of the foot [31]. Electrical stimulation devices are devices that attempt to provide a substitute neurological stimulation to the muscle tissue, which has lost its native neurological connection. Results of this treatment vary considerably and can be highly dependent on the cause of foot drop, and the time since foot drop was acquired. There are a variety of nuances to how electrical stimulation can be timed to a patients gait and appropriately triggered from

FIGURE 31.6 (A) Illustrations of a normal gait cycle; (B) ankle angle, (C) resistive and active joint power, and (D) muscle activity during the gait cycle with the solid line representing normal biomechanics and the dotted red line representing biomechanics during foot drop [32].

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various wearable sensors [32]. Neurological surgery or repair is a more likely treatment to recent foot drop acquired due to nerve injury, and is an attempt to restore the native neurological channel or provide a close substitute to enable voluntary dorsiflexion [31,33]. Tendon transfer procedures are intended to surgically recruit related muscle activity (via the posterior tibial tendon) to the line of action of the dorsiflexors [34].

31.7

Tarsal tunnel syndrome

31.7.1 Pathology Tarsal tunnel syndrome is a compression, or entrapment neuropathy of at least a portion or branch of the posterior tibial nerve. The nerve itself is located medial along the foot, and lies near the posterior tibial artery and vein, in a tunnel made of the calcaneus and fibrous fascia, under the flexor retinaculum. At this location, the nerve branches into the medial and lateral plantar nerves, and the medical calcaneal nerve. Therefore compression of this nerve or its branches, can result in paresthesia, dysesthesia, and hyperesthesia of the plantar foot or medial ankle—as these nerves provide for sensory and motor fiber innervation of the plantar foot and medial calcaneus [35 37]. Nerve potentials are affected by this syndrome (Fig. 31.7).

31.7.2 Impact on biomechanics This syndrome can include uni- or bilateral cramping and pain in the medial longitudinal arch of the foot. More directly, this syndrome may either prevent or incur weakness in abduction, adduction, flexion, and extension of the hallux. These effects are likely to occur later in the day after prolonged standing or walking [35,37].

31.7.3 Treatment A range of treatments may provide relief and recovery, including anti-inflammatory medication, positional control of the foot using custom orthotics or braces, night splits or a boot walker, or various stretching and strengthening exercises. Surgical intervention is considered following failure of conservative treatment, and particularly if the location of entrapment can be localized [35,37].

0.5 µV

1 µV 1 ms

1 ms

Left

Right

Medial Plantar

Lateral Plantar

FIGURE 31.7 Medial and lateral plantar nerve responses to a stimulus, compared between affected (right) and asymptomatic (left) side. The nerve potentials are significantly prolonged in the affected medial and lateral plantar nerve.

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FIGURE 31.8 The branching and division of the plantar nerves, note the fusiform bulge (Morton’s neuroma) in the nerve between the third and fourth digits.

31.8

Morton’s neuroma

31.8.1 Pathology Morton’s neuroma is a degenerative enlargement of the third common digital branch of the medial plantar nerve (Fig. 31.8). The term “neuroma” has been identified as a misnomer as this condition is not associated with the growth of new structure but rather with the degeneration of existing structures, including demyelination of nerves at this location and the formation of collagenous Renaut bodies [38] and fibrosis [39]. Metatarsalgia is another term for this condition, representing the pain present in the forefoot [40]. The pathomechanics leading to this condition are hypothesized to be due to muscle imbalance or deformity that impact the shape of the foot and the distribution of load during gait. For example, imbalances that prevent the forefoot from medial loading and increasing the load on the lateral rays may lead to this condition [40]. Footwear, such as pointed shoes, is also speculated to be a contributing cause [39]. Long second and third metatarsals may also lead to abnormal loading of the lessor rays, contributing to or instigating this condition [40].

31.8.2 Impact on biomechanics Painful gait can impact individuals differently and lead to different self-mitigation strategies. This condition has been described as “walking on a piece of stone or pebble” and may also yield either numbness or pain. Some patients feel worse with closed toed shoes, some feel worse barefoot [40]. The acute pain associated with this condition makes it unclear if it is tolerable long enough to develop secondary biomechanical conditions.

31.8.3 Treatment Conservative treatment approaches for Morton’s neuroma include massage and manipulation, insoles, and orthotics. Wide toebox shoes may be sufficient to alleviate symptoms [39]. The use of anti-inflammatory or alcohol injections is also part of the progressive treatment course [41]. Surgical treatment options include either nerve decompression or removal, and are typically performed through a dorsal or plantar approach [42]. Success rates for surgical treatment with neurectomy have been reported as high [39]. Neuroma in the presence of other conditions, such as hallux valgus or hallux rigidus, may require surgery to address both conditions [40].

31.9

Charcot foot

31.9.1 Pathology Charcot foot is a structurally devastating disease pathway whose origins are poorly understood. It has been associated with the prevalence of diabetes and/or peripheral neuropathy—potentially in combination with an initiating traumatic event. This condition is usually associated with advanced age, and a history of several years post diagnosis of diabetes or peripheral neuropathy. This disease affects the entirety of the foot tissue structures—soft tissues, bones and joints; and typically occurs in the midfoot and involves midfoot degradation and collapse [43 46]. This disease pathway has been described in four stages: inflammation, fragmentation, coalescence, and consolidation [45]. The typical end-stage presentation of Charcot foot results in a “rocker-bottom” deformity (Fig. 31.9).

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FIGURE 31.9 (A) Clinical presentation and (B) lateral radiograph of the rocker-bottom deformity of end-stage Charcot foot [45].

Patients presenting with early stages of this disease can be easily confounded with other lower extremity ailments as the outward symptoms are not unique. Imaging is a better determinant in the early stages of Charcot foot [44,45].

31.9.2 Impact on biomechanics Due to a high likelihood of concurrent peripheral neuropathy, Charcot foot can be painless, which can both minimize impact on gait biomechanics and increase disease severity and speed of progression. Associated features of this disease which impact biomechanics include an equinus contracture of the Achilles tendon; inflammation; and instability as the medial and lateral columns of the foot are compromised [43 45]. Due to potential patient comorbidities, normal gait biomechanics may already have been impacted in this population, preceding the onset of Charcot foot.

31.9.3 Treatment Offloading of the foot, to allow for the disease pathway to run its course with minimal change to the shape of the foot, is a “treatment” if this condition is caught early. This may be achieved through the use of custom casts (total contact casting), bracing, or even external fixation [44,45]. Cast walkers and knee scooters may also achieve the desired offloading, but have be associated with concerns of reliable use by patients [44]. Surgical treatment is also possible, but results can vary considerably based on disease stage and in this population carries high infection and necrosis risk [44]. Of the many injuries and diseases noted in this chapter, Charcot foot is an unfortunately impactful combination of devastating to foot function, difficult to predict/detect, and challenging to treat.

31.10 Charcot-Marie-Tooth disease 31.10.1 Pathology Charcot-Marie-Tooth (CMT) disease is defined by the National Institute of Neurological Disorders and Stroke as a hereditary motor and sensory neuropathy that damages peripheral nerves [47]. The musculoskeletal sequela of CMT include muscle weakness and atrophy, sensory loss, reduced tendon reflexes, and cavovarus foot deformity (a common deformity associated with CMT) [48 50]. Time is a significant factor as this a progressive disease.

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CMT was first described (and thus named) after French neurologists Jean Martin Charcot and Pierre Marie and British neurologist Howard Henry Tooth. This disorder has several “types” based on the presentation of genetic mutations. CMT is estimated to affect 126,000 individuals in the United States, and 2.6 million individuals worldwide [47]. CMT is the most common inherited neuropathy [51]. It should be noted that due to the diversity of genotypes that express the CMT phenotype, the extent, speed of development, and characteristics of CMT may vary significantly between individuals. This paradigm has caused substantial shifts in the diagnosis and targeted approaches to CMT with the rapidly increasing knowledge of human diseasecausing gene mutations. Mathis et al. surveyed over 100 CMT specialists worldwide and found majority consensus on updating the classification scheme of CMT based on advances in genetic identification of this disease.

31.10.2 Impact on biomechanics—pediatric and young adult A within-disease study by Pogemiller et al. evaluated differences in 3D kinematic and kinetic barefoot gait patterns in subjects with CMT Type 1 (n 5 20) and Type 2 (n 5 10) with an average age of participants in the mid-teens. They note overall findings of greater gait deviations in Type 2 versus Type 1 along with poorer disease severity scores in Type 2 [52]. An age matched, healthy versus CMT study in children by Ounpuu et al. evaluated ankle function differences. The CMT population (n 5 33) and healthy controls (n 5 21) populations had an average age of 12. The most common impairments found in the CMT population included ankle plantarflexor weakness and contracture, cavovarus foot deformity, and forefoot adductus. The observed ankle weakness commonly resulted in delayed peak dorsiflexion in terminal stance. The study suggested that this finding may be an early signature of progressive disorder [53]. Normal and CMT ankle kinematics and kinetics while walking up steps can be contrasted (Fig. 31.10).

31.10.3 Impact on biomechanics A survey of 407 respondents with CMT, aged 52.3 (SD 15.1) yielded responses of foot and ankle weakness (99.7%), impaired balance (98.6%), and mobility limitations (97.5%) [55]. A clinical evaluation of 98 CMT subjects performed an initial screening to identify those who did not have a primary gait defect (ankle ROM ratio and positive ankle work in stance phase) yielded 21 “normal like walkers” or CMT-NLW, this sub-population was contrasted against 31 healthy subjects who performed step ascending and descending tasks. The results of this study showed that despite “normal” level ground walking behavior, distal and proximal muscle weakness was apparent during ascent with some distal muscle weakness present in decent as well [54]. This indicates that even in high function level-ground walkers, impact on muscle impact due to the CMT disease process is present.

31.10.4 Treatment Treatments for this condition range are considerably based on age and disease severity. The spectrum of treatment ranges from physical or occupational therapy which can assist with management, orthotics, and splits may aid in

FIGURE 31.10 Mean profiles of the ankle kinematics and kinetics during step ascension in healthy subjects (thin gray line) versus CMT subjects (thick black line). Vertical lines represent phases: weight acceptance (WA), pull up (PU), forward continuance (FC), and swing (SW). The gray band represents normality range [54].

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correcting or compensating for mild deformities and surgical intervention for more severe deformities are typical [47,53]. As this is a progressive neurological disorder, musculoskeletal interventions can mitigate the effects but not address the underlying cause.

31.11 Friedreich’s ataxia 31.11.1 Background and pathology Friederichs’s ataxia (FRDA) is a multisystemic degenerative disease that first presents around adolescence and is caused by an autosomal recessive GAA triplet repeat expansion of the frataxin (FXN) gene [56,57]. It is the most common form of autosomal recessive ataxia, affecting 1 in 50,000 Caucasians [57]. Clinical presentation varies broadly, but common phenotypes include loss of lower limb reflexes, gait ataxia, sensory loss, and dysarthria. Scoliosis, pes cavus, and talipes equinovarus are common musculoskeletal deformities [56]. The frequency of certain clinical signs is positively correlated to the GAA expansion size or the disease duration (or a combination of both factors) [58]. The number of GAA repeats and duration of disease are both factors in the frequency of extensor plantar response, lower limb muscle weakness, and amyotrophy [58]. Scoliosis and pes cavus frequencies are influenced strongly by increased number of GAA repeats [58]. Due to the sensory deficits and cerebellar dysfunction, people inflicted with FRDA have difficulty maintaining balance; therefore, the ataxia gait is characterized by stumbling and unsteadiness and most patients are wheel-chair bound by about 10 15 years after disease onset [56 58].

31.11.2 Gait analysis It is well-known in the field that the gait measures vary broadly between FRDA patients, and thus investigations must be thorough and accurate. There are surprisingly few studies that conduct objective gait analysis and those few only focus on spatiotemporal parameters or only adult populations [57]. Spatiotemporal gait parameters are not indicative of true deficits in stability but rather a measure of the compensatory strategies to maintain balance [59]. Vasco et. al. investigates the changes in gait analysis over one year in a group of children and adolescents with FRDA. The results showed that when compared to healthy controls, subjects with FRDA showed slower walking speed, shorter and wider steps, longer stride duration, pronounced lateral displacement of center of mass, increased double stance time, and a decreased degree of peak plantarflexion [57,60]. The GAA1 (the smaller GAA repeat) repeat size has a negative correlation with the percentage of gait cycle in double support and a positive correlation with speed and the percentage of gait cycle in swing [59]. There are deviations in the hip, knee, and ankle angles from a normal gait cycle (Fig. 31.11).

31.11.3 Musculoskeletal effects Weakness and wasting are witnessed more often in the lower limb as opposed to the upper limb. 30% of cases have mild weakness in the lower limb, and 28% have severe weakness. Mild weakness and severe weakness in the upper limb are found in 45% and 2% of cases, respectively [61]. In typical cases of FRDA, uncomfortable and painful spasticity is prevalent, lower limb reflexes are absent, and muscle tone is reduced [61]. Several studies have witnessed some degrees of scoliosis in a large proportion of subjects, with most seeing scoliosis in more than 75% of cases [61]. Foot deformities are largely unaddressed in FRDA studies, though talipes equinovarus and pes cavus are observed, with pes cavus being the focus of the literature even though talipes equinovarus is more common in clinical cases [61]. Foot deformities can have a significant impact on the quality of life due to their implications for mobility or wheelchair transfers and positioning [61].

31.11.4 Clinical treatment There are no interventions that can halt the progression of the disease, so treatment mainly focuses on multidisciplinary management of symptoms. Some medications that change ion channel function and/or the physiology of the cerebellum can help treat some symptoms of the ataxia, but it is unclear whether these have the ability to modify the disease in any capacity. Riluzole is an example of a pharmaceutical symptomatic treatment that modulates SK channels to help regulate the firing patterns of neurons, and it improves speech and gait [62]. Physiotherapy involving passive stretching and aerobic exercise can help with balance, flexibility, strength, and coordination and temporarily help with spasticity. An orthotic device, splint, or cast might be necessary to provide more long-term muscle lengthening [56]. Physiotherapy treatments work to counteract the functional effects of FRDA by altering physiological processes, which may help

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FIGURE 31.11 The ankle, hip, and knee rotations during a complete gait cycle in healthy controls (solid line) compared to subjects with FRDA (dotted: baseline, dashed: during a follow-up after 12 months). Phases of the gait cycle are indicated by arrows: foot landing/start of gait cycle (dark gray) and the beginning of swing (white arrows). The large light gray arrow indicates progressive deterioration of the gait pattern in patients with FRDA, resulting in furthering extension of the ankle and knee [57].

delay, maintain, or improve functional degeneration [56]. Surgical intervention occurs in patients who have scoliosis (common with FRDA) that affects sitting and head control or the spinal curve is near 50%. There are several preclinical studies on rodents investigating disease-modifying treatments for FRDA that involve administering wild-type FXN through bone marrow transplants or injection of FXN-expressing adeno-associated virus [62].

31.12 Poliomyelitis 31.12.1 Pathology Poliomyelitis is the paralytic disease caused by the poliovirus. According to the CDC, most people infected with the virus will have no visible symptoms, 25% of people will develop flu-like symptoms, and a small portion of people will develop brain and spinal involvements which may range in severity from parathesia to paralysis and meningitis. Postpolio syndrome (PSS) is a sequalae of neuromuscular symptoms which may affect polio survivors years after the initial illness. These symptoms include weakness and joint pain [63].

31.12.2 Current status According to the World Health Organization, there were only 33 reported cases of polio in the world in 2018 [64]. Despite this good news and the reduced likelihood of future patients suffering from musculoskeletal pathologies due to poliomyelitis, there is a living population with PSS who require care and treatment.

31.12.3 Impact on biomechanics A Swedish study of 842 polio patients with a diagnosis of PSS included the measurements of patient characteristics, gait, and strength. Leg involvement was common, with only 3% of this population not having an effected leg versus 47% of this population not having arm involvement. Approximately 57% of this population did not use walking aids, of those that did the most common were crutches or canes (B26%). This population walked 71% slower than the selfselected walking speed of the general population. The dorsiflexion and plantarflexion strength of this population was

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50% 65% compared to the general population, and B28%/B8% of these patients were unable to generate a dorsiflexion/plantarflexion result, respectively [65].

31.12.4 Treatment The treatment (via prevention) for polio itself includes the inactivated poliovirus vaccine and oral poliovirus vaccine. For those who have severe cases of polio and who recover to experience PPS, physical therapy may restore some function, but many PPS patients utilize a variety of assistive or walking aids such as ankle-foot orthoses or knee-ankle-foot orthoses to manage their muscular weakness [66,67].

31.13 Areas of Future Research Despite the large number of individuals with neurological conditions that affect the foot and ankle, there has been limited research in this area. The often-heterogeneous effects of these conditions make studying them challenging, however improving our understanding of the changes in neuromuscular control that result from these conditions at the patientspecific level via in vivo measurements and computational simulation may help lead to the design of improved interventions.

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[18] Ruzbarsky JJ, Scher D, Dodwell E. Toe walking: causes, epidemiology, assessment, and treatment Curr Opin Pediatr 2016;28(1):40 6[cited 2022 Apr 1]. Available from: https://pubmed.ncbi.nlm.nih.gov/26709689/. [19] Oetgen ME, Peden S. Idiopathic toe walking J Am Acad Orthop Surg 2012;20(5):292 300[cited 2022 Apr 1]. Available from: https://pubmed. ncbi.nlm.nih.gov/22553101/. [20] Morozova OM, Chang TF, Brown ME. Toe walking: when do we need to worry? Curr Probl Pediatr Adolesc Health Care 2017;47(7):156 60 [cited 2022 Apr 1]. Available from: https://pubmed.ncbi.nlm.nih.gov/28716514/. [21] De Oliveira V, Arrebola L, De Oliveira P, Yi L. Investigation of muscle strength, motor coordination and balance in children with idiopathic toe walking: a case-control study Dev Neurorehabil 2021;24(8):540 6[cited 2022 Apr 1]. Available from: https://pubmed.ncbi.nlm.nih.gov/33759692/. [22] Alvarez C, De Vera M, Beauchamp R, Ward V, Black A. Classification of idiopathic toe walking based on gait analysis: development and application of the ITW severity classification. Gait Posture 2007;26(3):428 35. [23] Hirsch G, Wagner B, Hirsch G, Lindgren’s Barnsjukhus A. The natural history of idiopathic toe-walking: a long-term follow-up of fourteen conservatively treated children. [24] Eastwood DM, Menelaus MB, Dickens DRV, Broughton NS, Cole WG. Idiopathic toe-walking: does treatment alter the natural history? J Pediatr Orthop B 2000;9(1):47 9[cited 2022 Jun 2]. Available from: https://pubmed.ncbi.nlm.nih.gov/10647110/. [25] Fox AE, Deakin S, Pettigrew G, Paton R. Serial casting in the treatment of idiopathic toe-walkers and review of the literature Acta Orthop Belg 2006;72(6):722 30[cited 2022 Jun 2]. Available from: https://pubmed.ncbi.nlm.nih.gov/17260610/. [26] Stricker SJ, Angulo JC. Idiopathic toe walking: a comparison of treatment methods J Pediatr Orthop 1998;18(3):289 93[cited 2022 Jun 2]. Available from: https://pubmed.ncbi.nlm.nih.gov/9600550/. [27] Wrobel JS, Najafi B, Wrobel J. Diabetic foot biomechanics and gait dysfunction J Diabetes Sci Technol 2010;[cited 2021 Jun 1]. Available from: http://www.journalofdst.org. [28] Penman ID, Ralston SH, Strachan MWJ, Hobson R, editors. Davidson’s principles and practice of medicine. 24th ed. Elsevier; 2022. [29] Barrell K, Smith AG. Peripheral neuropathy Med Clin North Am 2019;383 97[cited 2021 May 28]. Available from: https://doi.org/10.1016/j. mcna.2018.10.006. [30] Fernando M, Crowther R, Lazzarini P, Sangla K, Cunningham M, Buttner P, et al. Biomechanical characteristics of peripheral diabetic neuropathy: a systematic review and meta-analysis of findings from the gait cycle, muscle activity and dynamic barefoot plantar pressure Clin Biomech 2013;831 45[cited 2021 May 28]. Available from: https://reader.elsevier.com/reader/sd/pii/S0268003313001897? token 5 4E8104F655EF43DBDC89A2F3C188A4EF93AA3CB6E108B0D252CA34ECF38F3E834562046447AEB6522A369E03F70FD2CC&originRegion 5 us-east-1&originCreation 5 20210528174936. [31] Carolus AE, Becker M, Cuny J, Smektala R, Schmieder K, Brenke C. The Interdisciplinary management of foot drop Dtsch Arztebl Int 2019;116(20):347[cited 2022 May 20]. Available from: https://pmc/articles/PMC6637663/. [32] Gil-Castillo J, Alnajjar F, Koutsou A, Torricelli D, Moreno JC. Advances in neuroprosthetic management of foot drop: a review. [cited 2022 May 20]. Available from: https://doi.org/10.1186/s12984-020-00668-4. [33] Sahu R, Garg RK, Malhotra S, Lalla R. Spastic foot-drop as an isolated manifestation of neurocysticercosis. [34] Schweitzer KM, Jones CP. Tendon transfers for the drop foot. Foot Ankle Clin 2014;19(1):65 71. [35] McSweeney SC, Cichero M. Tarsal tunnel syndrome—a narrative literature review. Foot 2015;25(4):244 50. [36] Nelson SC. Tarsal tunnel syndrome. Clin Podiatr Med Surg 2021;38(2):131 41. [37] Ahmad M, Tsang K, Mackenney PJ, Adedapo AO. Tarsal tunnel syndrome: a literature review. Foot Ankle Surg 2012;18(3):149 52. [38] Wr A. Morton’s neuroma Clin Podiatr Med Surg 2010;27(4):535 45[cited 2021 Aug 13]. Available from: https://pubmed.ncbi.nlm.nih.gov/20934103/. [39] Bhatia M, Thomson L. Morton’s neuroma—current concepts review J Clin Orthop Trauma 2020;11(3):406[cited 2022 Apr 1]. Available from: https://pmc/articles/PMC7211826/. [40] Gougoulias N, Lampridis V, Sakellariou A. Morton’s interdigital neuroma: instructional review EFORT Open Rev 2019;4(1):14[cited 2022 Apr 1]. Available from: https://pmc/articles/PMC6362341/. [41] Santiago FR, Mun˜oz PT, Pryest P, Martı´nez AM, Olleta NP. Role of imaging methods in diagnosis and treatment of Morton’s neuroma World J Radiol 2018;10(9):91 9Sep 28[cited 2022 Apr 1]. Available from: http://www.ncbi.nlm.nih.gov/pubmed/30310543. [42] Valisena S, Petri GJ, Ferrero A. Treatment of Morton’s neuroma: a systematic review Foot Ankle Surg 2018;24(4):271 81[cited 2021 Aug 13]. Available from: https://pubmed.ncbi.nlm.nih.gov/29409240/. [43] Rogers LC, Frykberg RG. The Charcot foot. Medical Clinics of North America, Vol. 97. Elsevier; 2013. p. 847 56. [44] Schmidt BM. Clinical insights into Charcot foot. Best Pract Res Clin Rheumatol 2020;34(3):101563. [45] Rosskopf AB, Loupatatzis C, Pfirrmann CWA, Bo¨ni T, Berli MC. The Charcot foot: a pictorial review Insights Imaging 2019;10(1)[cited 2022 May 20]. Available from: https://pmc/articles/PMC6682845/. [46] Trieb K. The Charcot foot: pathophysiology, diagnosis and classification Bone Jt J 2016;98-B(9):1155 9[cited 2022 May 20]. Available from: https://online.boneandjoint.org.uk/doi/abs/10.1302/0301-620X.98B9.37038. [47] Charcot-Marie-Tooth Disease Fact Sheet | National Institute of Neurological Disorders and Stroke. [cited 2021 Jan 8]. Available from: https:// www.ninds.nih.gov/Disorders/Patient-caregiver-education/Fact-sheets/Charcot-Marie-Tooth-Disease-Fact-Sheet. [48] Bird TD. Charcot-Marie-Tooth (CMT) hereditary neuropathy overview. GeneReviews. University of Washington, Seattle; 1993[cited 2021 Jan 15]. Available from: http://www.ncbi.nlm.nih.gov/pubmed/20301532. [49] Holmes JR, Hansen ST. Foot and ankle manifestations of Charcot-Marie-Tooth disease Foot Ankle Foot Ankle 1993;476 86[cited 2021 Jan 8]. Available from: https://pubmed.ncbi.nlm.nih.gov/8253442/.

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[50] Klein CJ. Charcot-Marie-Tooth Disease and Other Hereditary Neuropathies Contin Lifelong Learn Neurol 2020;26(5):1224 56[cited 2021 Jan 15]. Available from: https://journals.lww.com/10.1212/CON.0000000000000927. [51] OMIM—online mendelian inheritance in man. [cited 2021 Jan 15]. Available from: https://www.omim.org/. ˜ unpuu S. Comparison of gait patterns and functional measures between Charcot-Marie-Tooth [52] Pogemiller K, Garibay E, Pierz K, Acsadi G, O disease type I and II in children to young adults. Gait Posture 2020;77:236 42. ˜ unpuu S, Garibay E, Solomito M, Bell K, Pierz K, Thomson J, et al. A comprehensive evaluation of the variation in ankle function during [53] O gait in children and youth with Charcot-Marie-Tooth disease Gait Posture 2013;38(4):900 6[cited 2021 Jan 15]. Available from: https:// pubmed.ncbi.nlm.nih.gov/23702343/. [54] Lencioni T, Piscosquito G, Rabuffetti M, Di Sipio E, Diverio M, Moroni I, et al. Electromyographic and biomechanical analysis of step negotiation in Charcot Marie Tooth subjects whose level walk is not impaired. Gait Posture 2018;62:497 504. [55] Johnson NE, Heatwole CR, Dilek N, Sowden J, Kirk CA, Shereff D, et al. Quality-of-life in Charcot-Marie-Tooth disease: the patient’s perspective. Neuromuscul Disord 2014;24(11):1018 23. [56] Cook A, Giunti P. Friedreich’s ataxia: clinical features, pathogenesis and management British medical bulletin, Vol. 124. Oxford University Press; 2017. p. 19 30[cited 2021 Mar 26]. Available from: https://pubmed.ncbi.nlm.nih.gov/29053830/. [57] Vasco G, Gazzellini S, Petrarca M, Lispi ML, Pisano A, Zazza M, et al. Functional and gait assessment in children and adolescents affected by Friedreich’s ataxia: a one-year longitudinal study, 2016. [58] Du¨rr A, Cossee M, Agid Y, Campuzano V, Mignard C, Penet C, et al. Clinical and genetic abnormalities in patients with Friedreich’s ataxia N Engl J Med 1996;335(16):1169 75[cited 2021 Apr 9]. Available from: http://www.nejm.org/doi/abs/10.1056/NEJM199610173351601. [59] Milne SC, Hocking DR, Georgiou-Karistianis N, Murphy A, Delatycki MB, Corben LA. Sensitivity of spatiotemporal gait parameters in measuring disease severity in friedreich ataxia Cerebellum 2014;13(6):677 88[cited 2021 Apr 12]. Available from: http://link.springer.com/ 10.1007/s12311-014-0583-2. [60] Serrao M, Pierelli F, Ranavolo A, Draicchio F, Conte C, Don R, et al. Gait pattern in inherited cerebellar ataxias. [61] Parkinson MH, Boesch S, Nachbauer W, Mariotti C, Giunti P. Clinical features of Friedreich’s ataxia: classical and atypical phenotypes. J Neurochem 2013;103 17. [62] Kwei KT, Kuo S-H. An overview of the current state and the future of ataxia treatments. [cited 2021 Apr 12]. Available from: https://doi.org/ 10.1016/j.ncl.2020.01.008. [63] What is polio?. [cited 2021 Mar 5]. Available from: https://www.cdc.gov/polio/what-is-polio/index.htm. [64] Poliomyelitis. [cited 2021 Mar 5]. Available from: https://www.who.int/news-room/fact-sheets/detail/poliomyelitis. [65] Vreede KS, Sunnerhagen KS. Characteristics of patients at first visit to a polio clinic in Sweden PLoS One 2016;11(3)[cited 2021 Mar 5]. Available from: https://pubmed.ncbi.nlm.nih.gov/26981623/. [66] Portnoy S, Schwartz I. Gait characteristics of post-poliomyelitis patients: standardization of quantitative data reporting. Ann Phys Rehabil Med 2013;56(7 8):527 41. [67] Santos Tavares Silva I, Sunnerhagen KS, Wille´n C, Ottenvall Hammar I. The extent of using mobility assistive devices can partly explain fatigue among persons with late effects of polio—a retrospective registry study in Sweden. BMC Neurol 2016;16(1).

Chapter 32

Chronic Foot and Ankle Injuries Danielle Torp1 and Luke Donovan2 1

Department of Athletic Training and Clinical Nutrition, University of Kentucky, Lexington, KY, United States, 2Department of Applied Physiology,

Health, and Clinical Sciences, University of North Carolina at Charlotte, Charlotte, NC, United States

Abstract Chronic foot and ankle injuries are relatively common and are associated with numerous patient and clinical impairments. Proper rehabilitation and treatment techniques are imperative to restore impairments associated with chronic injuries to promote long-term health of the involved tissue(s). However, owing to the complex and multi-faceted mechanism by which chronic injuries develop, implementing effective treatment strategies can be challenging. An impairment-based rehabilitation model has been described and supported within the literature to address factors contributing to injuries. Therefore, the primary purpose of this chapter is to describe common pathways that lead to chronic injury, provide an overview of the most common foot and ankle chronic conditions, and to provide treatment and rehabilitation strategies to restore impairments associated with the injury. First, a theoretical framework is presented to describe pathways that lead to the development of chronic injuries. Next, associated risk factors, pathomechanics, and general characteristics of several chronic injuries are outlined. Finally, strategies to improve patient and clinical impairments of these pathologies are described along with general recommendations for treatment according to an impairment-based rehabilitation model.

32.1

Introduction

Chronic injuries can be defined as biomechanical disruptions of tissue integrity, which result in pain and a biochemical response that lasts longer than the normal acute inflammatory response [1]. Prior to presenting information regarding specific chronic injuries, we first describe a theoretical model about their development (Fig. 32.1). Based on this model, chronic injuries are driven by two main pathways: microtraumatic and macrotraumatic (acute injury). Both pathways can lead to patient and clinical impairments, which present as a chronic injury. Patient impairments are items that the patient reports such as pain, physical dysfunction, kinesiophobia, while clinical impairments pertain to items that clinicians measure such as range of motion (ROM), strength, balance, and biomechanics during movement [2].

32.1.1 Chronic injury through microtrauma All tissues (bone, cartilage, ligament, muscle, and tendon) within the body experience continuous repetitive stresses or forces. The quantity of stress that the tissue experiences is largely dynamic and is dependent on the body’s specific movement or task. For example, the amount of stress applied to the body’s tissue during running will be greater than that of walking. Under normal conditions, the tissues under stress experience microtrauma and some degradation, which triggers a cascade of events to restore and strengthen the tissue. This stress/repair cycle is essential for allowing the body to adapt to various forces across different tasks and structural loads. In fact, this process is continuously occurring with all tissue without the person being conscious of the cycle. Each tissue has a theoretical stress threshold [3]. The stress threshold should not be confused with the tissue’s tensile strength threshold (attributed with macrotrauma) which will be discussed later in this chapter, but rather pertains to repetitive forces tissue endures that triggers the stress/repair cycle. If the applied stresses are sub threshold, the tissue will both be repaired and strengthened as part of the stressrepair cycle; however, if the stress threshold is repetitively exceeded, the tissue may continue to break down or degrade resulting in a chronic injury (Fig. 32.1). Stated differently, under some circumstances, insufficient repair of the tissue Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00038-X © 2023 Elsevier Inc. All rights reserved.

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Tissue Repaired and Strengthened to Adapt to Stress Stress Repair Cycle Risk Factors (Nonmodifiable and Modifiable)

Insufficient Tissue Repair Balance Strength Range of Motion

Patient and Clinical Impairments

Normal Tissue Stress (microtrauma)

Altered Biomechanics Neuromuscular Control Muscle Activation Proprioception

Chronic Injury

Kinesiophopia Function Macrotrauma (Acute Injury)

Insufficient Rehabilitation/ Tissue Repair

Pain

Adequate Rehabilitation/ Tissue Repair

FIGURE 32.1 A theoretical model of the development of chronic injuries. Under normal conditions, tissue is exposed to microtrauma and repaired through the stress repair cycle. During this process, the tissue adapts to the imposed demands so that the tissue can withstand greater future stresses (green boxes). However, the presence of risk factors (yellow box) may disrupt the normal stress repair cycle and result in either macrotrauma (acute injury) (red box) or insufficient tissue repair (black box) leading to chronic injury comprised of various patient and clinical impairments (funnel). Following an acute injury, the tissue may (1) heal properly and continue a normal stress repair cycle (green boxes) or (2) experience inadequate healing and develop into a chronic injury (black box). This model also depicts that not all acute and chronic injuries (dashed line) are derived from risk factors.

occurs when the continued stress exceeds the threshold of the tissue [3]. The causes of why in some instances the tissue continues to breakdown remains unclear; however, some evidence suggests that the presence of risk factors (non-modifiable and modifiable) may disrupt the normal tissue stress repair cycle by inducing stresses that exceed the tissue’s threshold or lowers the stress threshold. We will operationally define non-modifiable risk factors as factors that cannot be altered by the patient or clinician (age, sex, height, bony alignment, etc.). On the other hand, modifiable risk factors are factors that can be altered by the patient or clinician (ROM, strength, balance, biomechanics during movement, diet, weight, etc.). Presently, the mechanism by which these risk factors increase the likelihood of developing chronic injury (exceeds tissue’s stress threshold or lowers the stress threshold) is not completely clear, but it can be speculated that factors related to influencing force transmission through the body (weight, bony alignment, biomechanics, etc.) contribute to exceeding the tissue’s threshold. On the other hand, factors such as diet and age most likely lower the tissue’s stress threshold. In the context of injury, the presence of risk factors, regardless of non-modifiable versus modifiable, means that an individual has a statistically greater chance of developing a chronic injury or sustaining an acute (macrotrauma) injury. To be clear, not all individuals with risk factors develop or sustain an injury. In fact, microtraumatic events, despite the presence of risk factors, may never surpass the threshold of tissue degradation and does not elicit a pain response thus no complaint of an injury arises [4]. However, understanding the role of risk factors during the normal tissue stress repair cycle is beneficial for clinicians attempting to reduce the rate at which individuals get injured or implement rehabilitation strategies to prevent the reoccurrence of injury. The presence of risk factors can alter the stresses placed on a tissue, the inflammatory response, the stress threshold, or a combination of all, ultimately impeding proper adaptation to the imposed demands. Through the stress/repair cycle, the body will attempt to alter the tissue’s stress threshold to accommodate the increased demands; however, the body is not always successful, which results in injury (acute or chronic). The presence of risk factors can result in three separate pathways: (1) acute injury, (2) chronic injury, and (3) no injury (Fig. 32.1). In addition to the separate pathways,

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presence of injuries, whether acute or chronic, may be a risk factor for additional acute or chronic injuries. Unfortunately, owing to the relationship between risk factors, chronic injuries, and acute injuries, the independent cause of the injury is often obfuscated.

32.1.2 Chronic injury through macrotrauma Macrotrauma or acute injury is the disruption of tissue stemming from a sudden force that exceeds the tissue’s tensile strength threshold [5]. Like chronic injuries, the presence of risk factors may increase the likelihood of an individual sustaining an acute injury. The severity of an acute injury is often described using a grading system (I III), where the grade corresponds to the degree of tissue disruption. Most commonly grade III injuries describe a complete rupture or tear of the tissue, grade II injuries describe partial tearing of the tissue, while grade I describes small tears or stretching of the tissue. Regardless of grade, macrotrauma will trigger a healing response. Oftentimes, acute injuries properly heal without any lingering clinical- or patient- oriented impairments; however, macrotrauma may serve as another pathway for developing chronic injuries (Fig. 32.1). In some cases following macrotrauma, the tissue does not properly advance through the phases of healing resulting in chronic inflammation, tissue degradation, and/or pain (chronic injury). Presently, the designation of the injury being chronic is often at the discretion of the clinician and based on changes to signs and symptoms as there is not a definitive time frame or event that represents the transition in classification. Aside from improper healing, the presence of an acute injury can alter a person’s biomechanics; thus, changes the stresses placed on other structures. If the new stresses exceed the corresponding tissue’s stress threshold, excessive microtrauma, as previously described, may be induced resulting in chronic injury. Prolonged presence of chronic injury can lower the tensile strength threshold of the tissue causing it to be more susceptible to future macrotrauma. The mechanistic development of a chronic injury is not straightforward; therefore, it is important to classify chronic injuries based on the primary mechanism causing the injury or condition. For example, the mechanism of some chronic injuries may remain unclear as it is possible for patients to present with an injury, but appear to have no risk factors or previous acute injuries.

32.1.3 Impairment-based rehabilitation model for treating chronic injuries The rehabilitative process concerning chronic injuries requires an understanding of the chronic inflammation process and how microtraumatic events can lead to an insufficient repair process. Furthermore, there is no specific timeline to determine when acute inflammation transitions into chronic inflammation; thus, it is left to the clinician to determine which stage of tissue healing is present. One strategy to guide the rehabilitation process is by implementing an impairment-based rehabilitation model. Rehabilitation of chronic injuries is not only reliant on the understanding of the etiology of the injury, but also the assessment of the injury characteristics. Stated differently, clinicians must be capable of detecting impairments that contribute to the chronic pathology. Therefore, utilizing an impairment-based rehabilitation model that incorporates an assess/intervene/re-assess approach may be the best strategy of resolving the condition and re-occurrence of the condition. The assessment component of the model should not merely focus on the clinical impairments (ROM, strength, balance, etc.) but also be comprised of patient-oriented impairments (pain, function, kinesiophobia, etc.). Furthermore, the re-assessment component should determine (1) whether the clinical impairment improved and (2) whether the improved clinical component corresponded to improved patient-oriented outcomes. The impairment-based rehabilitation model was originally developed as an intervention strategy for treating patients with chronic ankle instability (CAI) [6]. Since then, the philosophy of assess/intervene/re-assess impairments has been described to treat patients with other conditions such as acute ankle sprains [7] and patellofemoral pain [8]. Many chronic conditions share similar impairments when examining group means. For example, on average, patients with CAI have reduced dorsiflexion ROM when compared to individuals with no ankle sprain history [9]. Despite significant differences in dorsiflexion ROM between the two groups, this does not translate to every patient with CAI having reduced dorsiflexion ROM or vice-versa. For this reason, implementing a homogeneous rehabilitation plan for a condition that is comprised of heterogeneous impairments, is not practical. Incorporating an impairmentbased approach for treating patients with a chronic condition allows for the implementation of interventions or therapeutic exercises that are individualized to the patient’s specific deficiencies. Within this chapter, an overview of the etiology of chronic foot and ankle pathologies and how an impairment-based rehabilitation paradigm can be used to treat these injuries is described. Prior to presenting individual pathologies, it is important to provide an overview of the role of patient-oriented outcomes throughout the rehabilitation process.

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32.1.4 Role of patient-oriented outcomes Patient-oriented outcomes, or patient-reported outcome measures (PROMs), are questionnaires designed to quantify a patient’s subjective perception of their health condition (i.e., quality of life) and physical, psychological, or perceptual symptoms (i.e., dysfunction, depression, pain, etc.) including severity and frequency, which are undetectable to the clinician [10]. Preferably, these questionnaires are receptive to change, indicative of patient’s perception, and clinically reliable [11]. These instruments serve very distinct functions and commonly have a specific focus such as some PROM instruments are used to detect health-related quality of life, while others are region or anatomically specific [11]. For example, the 36-Item Short Form survey (SF-36) is used to measure generic quality of life pertaining to mental and physical health, while the Foot and Ankle Ability Measure is used to quantify perceived physical dysfunction associated with a foot or ankle pathology. Moreover, the Identification of Functional Ankle Instability [12] is a pathology-specific questionnaire created to detect specific criteria associated with functional ankle instability. Given the wide range of outcome measurements in the literature for quality of life, overall health, and foot and ankle specific pathologies, it is encouraged to incorporate both health-related quality of life, global, and joint specific questionnaires to gain an encompassing understanding of the treatment impact. Use of these instruments at various times during the rehabilitation process can provide useful information during each step of the assess/intervene/re-assess impairment-based model.

32.2

Chronic ankle instability

CAI is a musculoskeletal condition characterized by feelings of instability, repetitive episodes of the ankle “giving way”, and/or recurrent sprains, and persisting symptoms of pain, weakness, limited ROM, or functional deficiencies, which carry on greater than 1 year following the primary injury. CAI is unlike other chronic conditions as its onset follows inadequate recovery from an acute episode of ligamentous injury. Using our previously described model (Fig. 32.1), CAI is an example of a condition that originated from a macrotraumatic event that resulted in a chronic injury. Not all ankle sprains result in CAI as adequate tissue healing and rehabilitation may occur resulting in no lingering impairments (Fig. 32.1); these individuals are characterized as copers [13]. In the following sections we will discuss the pathophysiology and pathomechanics which cultivate after the initial lateral ankle sprain is sustained.

32.2.1 Anatomy overview The important bony structures involved with lateral ankle sprains include the tibia, fibula, talus, and calcaneus. The ankle joint complexes’ primary articulations include the distal tibiofibular syndesmosis, the talocrural joint, and the subtalar joint. These joints allow for weight bearing forces to transmit across the foot and lower leg. Dorsiflexion and plantarflexion occur at the talocrural joint, while the subtalar joint allows for inversion and eversion. It is important to note, during functional activities the ankle joint complex does not function in isolated movements of single joints. Therefore, it is unlikely for injuries to occur in a single plane of motion (i.e., lateral ankle sprains occur during inversion and internal rotation). Ligamentous support (Fig. 32.2) for the lateral ankle is provided by the anterior talofibular ligament (ATFL), posterior talofibular ligament (PTFL), the calcaneofibular ligament (CFL), and the anterior and posterior inferior tibiofibular ligaments. The AFTL resists ankle inversion and talar internal rotation and is taut during plantarflexion. The PFTL limits posterior displacement of the talus. The CFL is the primary inversion resistor during mid-ROM of the talocrural joint [14,15].

32.2.2 Etiology 32.2.2.1 Mechanism of injury and pathomechanics Disruption of the lateral ligaments occurs when there is a forceful inversion and internal rotation at initial contact with the ground or with another object (Fig. 32.3) [16]. The most commonly injured ligament during an inversion ankle sprain is the ATFL, closely followed by the CFL. The structural damage to the ligaments initiated during an initial ankle sprain and the mechanism in which they occur are similar to that which transpire during recurrent ankle sprains. Following an ankle sprain, between 30% and 70% of individuals present with several patient-, clinical-, and laboratory-impairments which contribute to the development of CAI [17]. The specific impairments fall into one of three main categories: pathomechanical, sensory-perceptual, and motor-behavioral [18]. It is important to note not every

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FIGURE 32.2 Lateral ankle ligamentous support. From Elsevier publication: https://www.sciencedirect.com/topics/medicine-and-dentistry/anklejoint. Reproduced with permission from Elsevier, Drake et al. 2015.

FIGURE 32.3 Video-captured lateral ankle sprain recreated as a virtual environment (bottom right). Analysis of the video showed the mechanism of injury of a lateral ankle sprain to be excessive ankle inversion accompanied with tibial internal rotation. From Fong, Daniel TP, et al. A lateral ankle sprain during a lateral backward step in badminton: a case report of a televised injury incident. J Sport Health Sci 2021.. Reproduced with permission from Elsevier.

individual will present with every impairment, and as such, an assess/intervene/re-assess model should be used to target specific deficiencies present [6,8]. Pathomechanical impairments are considered structural deviations contributing to ankle joint dysfunction and chronicity, which arise secondary to the index ankle sprain and may affect the ankle joint and/or surrounding tissues [18]. Specific pathomechanical insufficiencies include pathological laxity, arthrokinematic and osteokinematic restrictions, and primary and secondary tissue adaptations. Sensory-perceptual impairments can be physiological, psychophysiological, or psychosocial components which the individual perceives or experiences regarding their injury, body, or self [18]. Such insufficiencies include pain, diminished somatosensation, perceived instability,

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kinesiophobia, lower self-reported function, and reduced health-related quality of life. Motor-behavioral insufficiencies include altered reflexes, neuromuscular inhibition, muscle weakness, balance deficits, altered movement patterns, and reduce physical activity. These impairments encompass deviations or deficiencies during movements an individual performs, or choses to avoid, including biomechanics and muscle activation patterns [18]. A deficient sensorimotor system has been well documented in patients with CAI. The sensorimotor disruption has been previously described as a continuous spectrum ranging from sensory deficits (proprioception) [19] to motor deficits (alterations in movement patterns) [20], with numerous other deficiencies (e.g., motor neuron pool excitability, reflex reactions, strength, postural control) between the ends of the spectrum [21]. Pertaining to alterations to biomechanics during movement, typically, individuals with CAI present with a more inverted foot position at initial contact [22] and spend more time on the lateral aspect of their foot during the stance phase of gait [23,24]. This anomalous position is thought to place the ankle in a position similar to the mechanism of injury for an inversion ankle sprain, thus leading to further recurrent sprains.

32.2.3 Clinical impairments Individuals with CAI often present with deficits that can be categorized within four broad domains [6]. The impairments in each of these domains may not be present in every individual with CAI; however, previous research supports group differences exist across these domains between those with and without CAI. Therefore, it is important to assess all areas in each domain to ensure the presence or lack of abnormalities are known prior to treatment. The exact mechanism for developing CAI is not completely understood; however, there is sufficient evidence to support the presence of these impairment domains in individuals with CAI.

32.2.3.1 Range of motion Dorsiflexion ROM is the most common ankle motion to present with deficiencies. It is believed decreased dorsiflexion ROM is attributed to an anterior positional fault of the talus and fibula which occurs following a lateral ankle sprain [25,26]. The anteriorly displaced talus relative to the tibia is thought to contribute to arthrokinematic and osteokinematic restrictions associated with CAI [27]. The decreased dorsiflexion ROM may cause a more plantarflexed position during gait, leaving it in a more open-packed and unstable position, which could predispose the joint to recurrent inversion injury.

32.2.3.2 Strength Strength deficits in all four of the major muscle groups that cross the ankle joint have been previously reported as a characteristic of CAI [19,28]. Weakness of the ankle evertors and dorsiflexors may contribute to the more inverted and plantarflexed position of the foot during gait. There have also been recent reports of proximal joint muscle weakness contributing to gait alterations at the ankle, therefore it is important to assess strength at the hip, knee, and ankle.

32.2.3.3 Balance Static and dynamic postural control have been found to be significantly affected in cases of CAI [29,30]. Diminished proprioceptive and neuromuscular control capabilities have also been found in individuals with CAI.

32.2.3.4 Functional activity Reduced performance during functional activities (e.g., walking, running, jumping, cutting, etc.) is a common complaint from individuals with CAI. It is likely the aforementioned alterations in ROM, strength, and balance all contribute to the decreased functional capabilities. As previously mentioned, CAI is associated with anomalous gait patterns which place the foot in a predisposing position for inversion injury while walking [31,32] and jogging [32]. Furthermore, it has been found this inverted ankle position is also seen during more functional tasks such as jumping.

32.2.4 Treatment 32.2.4.1 Acute management CAI is a condition which is thought to arise primarily from inadequate healing from an index inversion ankle sprain. Therefore, proper management and rehabilitation should be of importance to prevent chronicity. The most recent

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recommendations for early management of inversion ankle sprains to maximize healing and prevent onset of CAI have been summarized into 4 broad categories [33 36]. First, the application of ice, compression, and elevation in conjunction with non-steroidal anti-inflammatory is used to target pain and inflammation. Second, it is recommended to begin non-weight bearing motion and weight-bearing activities as soon as tolerated, rather than prolonged immobilization. Third, an external support such as ankle braces or tape should be used in the 12 months following an ankle sprain. Lastly, weight-bearing activities such as balance and coordination training should begin as soon as they can be tolerated.

32.2.4.2 On-going management Restoring adequate dorsiflexion ROM can be done with talar and fibular joint mobilizations [37]. Addressing arthrokinematic restrictions should be done prior to osteokinematic motions, as the former may disguise the outcomes of the latter. Ensuring flexibility of the triceps surae group can also aid in the restoring of dorsiflexion ROM. Strength training should be implemented for the weakness found during initial assessment and should be adjusted as progress and re-assessment indicates continuation or cessation of exercises. Both concentric and eccentric deficits should be addressed during the rehabilitative process. Balance training has been found to be an effective method for improving postural control, proprioception, and neuromuscular control [38]. Furthermore, improvements of patient-reported physical activity was most effectively accomplished through balance training [39]. While there are various protocols and methods for balance training, inclusion of both static and dynamic exercises are recommended considering both forms of postural control are deficient with CAI. The introduction to functional activity should occur during restoration of the deficits found in the previous three impairment domains. While ROM, strength, and balance impairments are being properly restored, the clinician should also focus on restoring proper walking gait mechanics. While no formal protocol exists for gait re-education, emphasis should be placed on abnormalities found during assessment of gait. Once proper walking gait mechanics, ROM, strength, and balance impairments have been addressed, progression into functional activities of running, jumping, and cutting should be implemented.

32.3

Plantar fasciitis

Plantar fasciitis is a musculoskeletal condition involving the plantar aponeurosis and is most commonly characterized by pain at the medial calcaneal tubercle during the first few steps following a prolonged period of non-weight bearing. It is important to note while the suffix “itis” typically refers to “inflammation of,” most cases of chronic plantar fasciitis do not exhibit inflammation; rather, the tissue undergoes chronic stretching, tearing, or degeneration at its attachment site. Referring back to our injury model (Fig. 32.1), plantar fasciitis can arise from microtrauma within the tissue and in the presence of risk factors lead to a macrotraumatic event or chronicity of improper tissue healing or degeneration [40].

32.3.1 Anatomical overview The central slip of the plantar fascia originates at the medial calcaneal tubercle and the medial and lateral slips emerge as the dense connective tissue fans out distally to its attachment on the metatarsophalangeal (MTP) joints of each digit. The plantar fascia has two purposes: (1) support and restore (via the windlass mechanism) the medial longitudinal arch and (2) act as a dynamic shock absorber of the foot [41]. The plantar fascia is composed of distal three slips which all arise from the proximal central slip.

32.3.2 Etiology 32.3.2.1 Mechanism of injury and pathomechanics During weight-bearing, the metatarsal heads fan out and the plantar aponeurosis is elongated. When the foot is moved into a toe-off position, the plantar aponeurosis is further stretched and increased tensile forces are present. During toeoff of both walking and running, the tensile forces on the plantar fascia are increased to upwards of 2x the body weight of the individual [42]. This increased loading can lead to microtrauma which initiates the inflammatory process if other structures of the foot (intrinsic and extrinsic foot muscles/tendons) cannot properly dissipate the forces across the entire structure. Under the normal stress repair cycle, the microtrauma within the plantar fascia tissue is repaired; however, in

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the presence of risk factors, the loading of the plantar fascia becomes excessive and the microtrauma to the area inhibits the repair process. Therefore, like many other chronic conditions, the tissue is fixed in a cycle of trauma and initial tissue repair without the ability for the inflammation process to resolve leading to chronic tissue degradation (Fig. 32.1). Specific biomechanical and structural risk factors place the plantar fascia under abnormal stresses include: pes planus, pes cavus, excessive hindfoot eversion or forefoot varus during the stance phase of gait, inflexibility of the longitudinal arch, and tightness of the triceps surae group [43,44]. Aside from the biomechanical risk factors, other modifiable risk factors such as training errors, repetitive exercise bouts on a hard surface, increased body mass index, and improper footwear have been identified. Regardless of which aforementioned risk factor or factors are present, the ability for the foot to dissipate forces is impaired and often the forces accumulate within the origination of the plantar fascia.

32.3.3 Clinical impairments 32.3.3.1 Range of motion Active and passive toe extension, in combination with dorsiflexion, will elicit pain. Restricted dorsiflexion ROM is considered the most important risk factor [45], and may be limited due to overall tightness of the triceps surae group.

32.3.3.2 Strength A single leg heel raise will identify any weakness or neuromuscular control deficits of the leg. Pain stemming from the plantar fascia may cause an altered gait which could progress to atrophy of the dorsiflexors and plantarflexors, as well as the intrinsic foot musculature. The perpetual cycle of pain and muscle weakness further eliminates the foot’s ability to act as a rigid to mobile to rigid system during gait, which ultimately causes the plantar fascia to be placed under greater stresses.

32.3.3.3 Functional activity The limited dorsiflexion ROM in individuals with plantar fasciitis will restrict movement during gait and interrupt the windlass mechanism [42], resulting in a compensatory mechanism by the hindfoot [46]. Furthermore, upon initial contact during gait, the foot is often in a more inverted and plantarflexed position. The increased frontal plane motion during gait creates an additional shear stress within the plantar fascia. Given the longitudinal orientation of the plantar fascia fibers, the overall structure may deteriorate as fibers of the plantar fascia begin to tear in the presence of the additional cross-sectional stress.

32.3.4 Treatment 32.3.4.1 Acute management Early pain management is required prior to addressing biomechanical adaptations. Early stages of plantar fasciitis require a state of rest and partial non-weight bearing to allow the repair process proper time to heal any damaged tissue. Once pain has subsided, focus should then be turned to other patient- and clinical impairments.

32.3.4.2 On-going management of clinical impairments Restoration of ankle dorsiflexion and first MTP sagittal plane motion is essential during the rehabilitative process. Regaining these motions will allow for proper function during gait. Regaining strength of the intrinsic and extrinsic foot musculature is important to provide stability to the foot and support the plantar fascia. Fully functioning intrinsic and extrinsic musculature improves dynamic stability of the foot and ankle joint complex, which dissipates excessive stressed being placed on static stability structures such as the plantar fascia. From a functional standpoint, supportive footwear, heel cup, or orthoses should be worn during weight bearing activities, especially for prolonged periods of time, to compensate or correct the biomechanical malalignments.

32.4

Tendinopathy (Achilles, peroneal, and posterior tibialis)

Tendinopathy (disease of a tendon) can manifest within the main structure of a tendon, at the insertion site (entheses), or in the structures surrounding the tendon (synovium/peritendon) [47,48]. Not only can tendinopathy present in various locations, but it evolves in three common forms: tendinitis, tendinosis, and paratenonitis [49]. It is important to briefly

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discuss the main differences between the conditions and to keep these definitions in mind when performing injury assessments as it will assist in making the appropriate distinction. Tendinitis is characterized by the presence of acute inflammation of a tendon resulting from micro-tears in the structure due to increased tensile forces that are too rapid or too great in magnitude, thus overloading the musculotendinous segment. Tendinitis is considered solely a clinical symptom, therefore should only be used when the clinical signs and symptoms of acute inflammation are present. Thus, most overuse or chronic pathologies are considered to be tendinosis or paratenonitis in nature [50]. Tendinosis is the degeneration of collagen and disorientation of tendon fibers, without the presence of clinical or histological signs of inflammation, resulting from chronic overuse or repetitive strain without sufficient time for the tendon to heal [48]. Lastly, paratenonitis refers to the inflammation of structures surrounding a tendon (peritendinitis, tenosynovitis, and tenovaginitis) at locations where tendons are subject to high friction. These structures may become inflamed due to chronic overuse and improper rest and healing between stresses. The distinction between these tendinopathies can be made with a proper understanding of injury development and presentation of symptoms. Regardless of which tendinopathy is present the goal remains the same: re-establish tensile strength (i.e., eccentric strengthening) and break the trauma cycle to allow the repair process to occur without disruption.

32.4.1 Anatomical overview The tendons of the medial and lateral gastrocnemius combine with the soleus to create the Achilles tendon which has a common, broad insertion on the posterior calcaneus. Its primary function is to plantarflex the ankle. The posterior tibialis originates on the posterior tibia and interosseous membrane, proceeds distal and medial, runs posterior to the medial malleolus before attaching on the navicular. The posterior tibialis provides dynamic stabilization to the medial longitudinal arch and actively plantarflexes and inverts the foot. The peroneal tendons reside in the lateral compartment of the leg, the longus originates on the proximal two thirds of the fibula while the brevis originates on the distal one third. They both run distal and posterior to the lateral malleolus. The peroneus longus tendon continues along the plantar aspect of the foot before attaching to the medial cuneiform and first metatarsal, while the peroneus brevis tendon attaches at the base of the fifth metatarsal. The peroneal tendons provide dynamic lateral ankle stability.

32.4.2 Etiology 32.4.2.1 Mechanism of injury and pathomechanics 32.4.2.1.1 Achilles tendon The repetitive shortening and lengthening of the Achilles tendon during the functional activities of running and jumping increase susceptibility for overuse and microtrauma (Fig. 32.1) [47]. The constant transition from concentric to eccentric loading threatens the ability of the musculotendinous unit to properly absorb forces. Inadequate healing in response to excessive tensile loading forces contributes to the chronic development of Achilles tendinopathy [51]. Biomechanical risk factors contributing to Achilles tendinopathy have been identified as pes cavus, forefoot varus, tibial varum, and gastrocnemius-soleus inflexibility. These anatomical malalignments subsequently cause excessive frontal plane motion of the hindfoot and an increased dorsiflexion during heel contact, ultimately increasing stress and development of tendinopathy [52]. Achilles tendinopathy can be differentiated as insertional or non-insertional, based on anatomical location of signs and symptoms. Furthermore, the type of tendinopathy most commonly found is tendinosis, but tendinitis and tenosynovitis are also possible. Considering the avascularity of the Achilles tendon 2 6 cm proximal to the insertion site, noninsertional Achilles tendinosis and tenosynovitis are likely in this area, whereas insertional Achilles tendinitis is more common due to the increased vascularity at the calcaneal attachment [53]. Distinguishing between insertional and noninsertional primarily require visual inspection to determine location of swelling or tendon enlargement, along with patient-reported location of pain. 32.4.2.1.2 Tibialis posterior tendon Tendinopathy occurring at the posterior tibialis is most often associated with tendinosis. Commonly referred to as posterior tibial tendon dysfunction (PTTD), the tendinosis pathology inhibits the tendon function in providing dynamic support to the medial longitudinal arch during weight bearing due to degradation of collagen bundles [54]. The development of posterior tibial tendinosis is attributed to both macrotraumatic and microtraumatic mechanisms [55,56]. Patient-reported pain can stem from inflammatory factors in the tendon or from gait deviations arising due to tendon

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inhibition [54]. PTTD is associated with reduced hindfoot eversion and impaired proprioceptive abilities when walking on uneven ground, and during stair ascent and descent [55]. Pes planus, hindfoot valgus, and forefoot abduction are common anatomical malalignments associated with tendinopathies affecting the posterior tibialis tendon [55,57]. 32.4.2.1.3 Peroneal tendon Several mechanisms have been identified as placing excessive mechanical stresses on the tendons which could lead to tendinopathy. A flexed first ray and forefoot valgus deformity provide a mechanical disadvantage to the peroneus longus during the stance phase of gait [58]. An anatomical hindfoot varus position places excessive tensile forces on the peroneal tendons as they work to evert the malposition. Peroneal tendinitis and tenosynovitis can manifest in either the peroneus longus, brevis, or in the shared synovial sheath posterior to the lateral malleolus [59]. Pain presents posterior and distal to the lateral malleolus and follows along the peroneal tendons. Excessive passive hindfoot inversion and ankle plantarflexion and resisted hindfoot eversion and ankle dorsiflexion will intensify pain [59].

32.4.3 Clinical impairments Tendinopathy conditions will present with painful active and passive motions, with or without the presence of crepitus. Prolonged periods of immobility or inactivity result in ankle joint stiffness which will reduce as motion and activity increase. A majority of tendinopathies are associated with weakness and inhibition in the corresponding muscles. Weakness in muscles associated with tendinopathy mainly include eccentric strength, but concentric and isometric strength is likely to also be reduced. Pain may be present during strength assessments. Careful assessment for biomechanical alterations should determine the influence of soft tissue restricting motion or structural alignment. For example, limited dorsiflexion ROM during ambulation and stair-ascent can be credited to triceps surae inflexibility or forefoot varus. Distinguishing which factor is influencing the anomalous gait will provide insight into treatment.

32.4.4 Treatment Increasing triceps surae flexibility will restore dorsiflexion ROM and reduce tension on the Achilles tendon or on the other involved tendons. Any arthrokinematic restrictions should also be treated with joint mobilizations to restore foot and ankle motion. There is an accepted consensus that eccentric strengthening is the most effective treatment for the restoration of pain and function from tendinopathy [60 62]. Eccentric strengthening in a controlled, therapeutic manor is able to restore proper collagen alignment and give the tendon the ability to withstand eccentric loads during functional activities. A 10-week eccentric strengthening program for PTTD saw significant reductions in clinical- and patient-oriented outcomes; however, ultrasonography comparisons did not show any morphological changes in the tissue [60]. Even though there were no visible changes to tissue, patients in this study maintained their improved pain and function 6 months after ceasing treatment [60]. Anomalous biomechanics which occur due to pre-existing structural anatomical deformities should be addressed with the use of orthotics. Biomechanical alterations arising from ROM deficits due to arthro- or osteokinematic restrictions should be appropriately addressed before introducing appropriate gait re-training.

32.5

Stress fractures (navicular, metatarsals)

32.5.1 Anatomical overview The medial and lateral portions of the navicular have vascular networks branching from the dorsalis pedis and posterior tibial arteries; however, the central third of the bone remains relatively avascular. The metatarsals are long bones which undergo torsion and compression forces during weight-bearing.

32.5.2 Etiology 32.5.2.1 Mechanism of injury and pathomechanics Stress fractures are the result of repetitive microtraumatic external forces outweighing a bones ability to dynamically respond via tissue remodeling (Fig. 32.1) [63,64]. Stated differently, the lack of adequate time between stressors does not allow the remodeling mechanism to maintain suitable bone strength. When bone strength falls below the “critical stress threshold,” a fracture can develop [63]. While the mechanism for stress fracture development is multifactorial

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and variant depending on the exact location, there have been several proposed factors generally associated with stress fractures. These include muscle function (i.e., fatigue, imbalance, or atrophy) [64,65], anatomical malalignments and biomechanical alterations [66,67]. The navicular is the anatomical link between the hindfoot and the forefoot further predisposing it to acquired stress from surrounding bones, thus increasing the susceptibility for stress injury (Fig. 32.4). During gait, ground reaction forces from the first and second metatarsals will transfer to the navicular. The central third of the navicular undergoes excessive sheer stress from unequal compressive forces from the first and second metatarsals heads [68]. A short first metatarsal and relatively long second metatarsal are anatomical characteristics commonly shared in navicular stress fractures [67], which likely further exacerbates the imbalanced force distribution on the navicular. Another theory postulates a restricted dorsiflexion and subtalar motion during gait will impinge the navicular resulting in repetitive microtrauma [66,68]. Stress fractures of the metatarsals are divided into low- and high-risk classifications based on anatomical location of fracture [69]. Low-risk stress fractures in the mid to distal portion of the second and third metatarsals are considered the most common with an incidence rate of 80% 90% [70] of all metatarsal stress fractures. Stress fractures in the first and fifth metatarsals are considered high risk and are rarer in occurrence. A majority of metatarsal stress fractures arise from high-risk activities, which involve excessive and prolonged stress loads such as running, jumping, dancing, gymnastics, and military service [64]. The relative immobility and load acceptance of ground reaction forces of the second and third metatarsals during weight-bearing likely increases risk of stress injury. Reduced dorsiflexion ROM and a forefoot varus deformity are associated with an increased risk of metatarsal stress fractures [71].

32.5.3 Clinical impairments The point of mechanical failure in the navicular is not always directly correlated to manifestation of patient-reported pain. Nonetheless, the presence of pain will likely produce an altered gait pattern in attempts of reducing stress on the tissue [72]. Point tenderness over the injured foot and increased pain with repetitive and prolonged weight-bearing activities are clinical signs of a stress fracture. There may be signs of inflammation (i.e., swelling, edema, ecchymosis,

FIGURE 32.4 Navicular stress fracture shown on sagittal slice of CT scan. “File:NavicularFracMark.png” by James Heilman, MD is licensed under CC BY-SA 4.0.

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etc.), but this will depend on the stage of injury. Active and passive ROM typically is not limited with stress fractures; however, forced end ROM that places stress on the injured bone will elicit pain and discomfort.

32.5.4 Treatment Stress fractures are conservatively treated with a prolonged period of non-weight bearing to allow bone remodeling and repair to adequately occur without disruption. Low-load and low-force strengthening can occur during this period of non-weight bearing in an attempt to maintain ROM. Strength exercises should consider the mechanical stress the tendon will place on bones. For example, strengthening of the posterior tibialis should be closely monitored so the tendon does not mechanically overload the healing navicular. Once bone healing has been radiographically confirmed (Fig. 32.5), progressive weight-bearing should begin with a focus on addressing any gait alterations that were identified as inducing the stress fracture. Other factors that may have contributed to the stress fracture, such as metabolic disorders, should also be evaluated [73].

32.6

Sesamoiditis

The term sesamoiditis typically refers to the inflammation or degradation of the flexor halluces tendons, or other surrounding tissue, which house the sesamoid bones.

32.6.1 Anatomical overview The two hallucal sesamoids are embedded within the medial and lateral flexor hallucis brevis (FHB) tendons and provide a mechanical advantage to the tendons. They serve to absorb and dissipate forces during weight bearing from the first metatarsal head, augment the gliding ability of the FHB tendons over the first MTP joint, and provide protection for the flexor hallucis longus tendon [74,75].

32.6.2 Etiology 32.6.2.1 Mechanism of injury and pathomechanics The mechanical advantage that the sesamoids provide to the FHB during normal gait and weight bearing is critical for overall foot mechanics [75]. During toe-off of gait, the sesamoids can transmit loads upwards of 300% of body weight. Understandably, the sesamoids and the surrounding tissues are susceptible to injury. Anatomical and biomechanical abnormalities will place excessive strain and stress on the sesamoids and surrounding tissue which may eventually lead to complaints of pain and swelling during weight-bearing activities. Such abnormalities include pes cavus, flexed first ray, gastrocnemius inflexibility leading to hindfoot inversion, and tibial internal rotation. The repetitive microtrauma occurring with these abnormalities leads to overuse and disrupted repair processes eventually leading to mechanical failure of the tissue. Active and passive motions involving the first MTP joint, especially when the ankle in already in full dorsiflexion, will be painful and restricted [76]. A flexed first ray during non-weight bearing is indicative of a shortened flexor halluces brevis tendons, which in turn will also reduce the extension motion of the MTP joint. The overextension of the plantar tissues will place excessive forces upon the sesamoids, eliciting pain. Sesamoiditis is considered a non-traumatic overuse condition, which is commonly associated with training errors excessively loading unconditioned tissues.

32.6.3 Clinical impairments Radiographs should be obtained to rule out any severe sesamoid injuries (e.g., acute or stress fractures, avascular necrosis, etc.). Diagnosis of sesamoiditis is typically confirmed following negative radiographs for the previously mentioned disorders. Restricted dorsiflexion and MTP joint extension are accompanied by pain during active and passive motions. Chronic pain and restricted motion could inhibit the intrinsic foot musculature, resulting in a general weakness and instability in the mid- and forefoot. Restricted dorsiflexion and MTP joint extension will cause abnormal gait mechanics, in addition to an altered push-off phase to avoid pain.

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FIGURE 32.5 Anterior-to-posterior (A) and oblique (B) radiographs of a third metatarsal stress fracture. Bottom images (C) and (D) correspond to radiographs repeated 6 months post injury. From Boukhemis K et al. Second metatarsal stress fractures. Oper Tech Orthop 2018;28(2):84 90. Reproduced with permission from Elsevier.

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32.6.4 Treatment Soft tissue mobilizations are needed to restore ROM of the first MTP joint, along with dorsiflexion at the ankle. Intrinsic foot strengthening should be included to provide stabilization through the mid- and forefoot and the medial longitudinal arch. Addressing any strength deficits of the extrinsic musculature, namely of the medial and posterior compartments of the leg. Biomechanical abnormalities need to be addressed to remove the cause of excessive stress being placed on the sesamoids. Anatomical variations can be corrected using custom orthoses designed to offload the sesamoid complex.

32.7

Retrocalcaneal bursitis

32.7.1 Anatomical overview The retrocalcaneal bursa is situated between the anteroinferior Achilles tendon and the posterosuperior calcaneal prominence [77,78]. The bursa is a thin, synovial fluid-filled sac tasked with reducing friction and acting as a cushion between the Achilles tendon and the calcaneus. The retrocalcaneal bursa lining is heterogeneous; the anterior portion in contact with the calcaneus is cartilaginous while the posterior portion in contact with the Achilles tendon is more tendinous, and the proximal lining is standard synovium [79].

32.7.2 Etiology 32.7.2.1 Mechanism of injury and pathomechanics The development of retrocalcaneal bursitis is almost exclusively attributed to a microtraumatic mechanism of mechanically induced inflammation associated with anatomical deformities allowing excessive coronal plane motion of the ankle [80]. However, the bursa can become inflamed and enlarged from a forceful direct blow to the area. An enlarged posterior prominence of the calcaneus reduces the space between the bone and the Achilles tendon, which impinges the retrocalcaneal bursa during dorsiflexion. Other anatomical and biomechanical anomalies which allow unnecessary coronal plane motion include a rigid flexed first metatarsal, pes cavus, and compensated hindfoot varus and forefoot valgus [80]. The unaddressed biomechanical compensations cause friction between the bursa and the calcaneus which will induce an inflammatory response in the bursal sac. The cumulative friction and pressure will eventually lead to a chronic inflammatory state, increasing the size of the bursa causing less space and further impingement. While the purpose of the bursa is to reduce friction and allow the Achilles tendon to glide over the calcaneus, abnormal mechanical stress will induce an inflammatory response and prolonged stress will result in degeneration or calcification of the fibro-cartilaginous bursa [78,81]. Histological examinations have also identified a buildup of synovial fluid and a hypertrophic inner sac lining [78,81].

32.7.3 Clinical impairments Patient-reported pain over the heel and visible inflammation proximal to the calcaneus and medial and lateral to the Achilles tendon. Active and passive dorsiflexion ROM will impinge the bursa and elicit a painful response. ROM may also be limited due to a compensated gait resulting in a shortened Achilles tendon.

32.7.4 Treatment The overarching goal in conservative treatment for retrocalcaneal bursa is to remove the exacerbating biomechanical factor(s). This approach works well for reducing irritation and reduction of fluid inside the bursa. Altering the anatomical malalignments and adverse mechanical loading patterns will reduce the friction placed on the bursa. Restoring triceps surae length will assist with ROM impairments. Chronic bursitis is commonly treated with non-steroidal antiinflammatory medications and possible corticosteroid injections. However, caution has been advised in treating bursitis with corticosteroid injection as it has been associated with an increased risk of Achilles tendon rupture [77].

32.8

Areas of future research for chronic foot and ankle injuries

It is important to note that many of the injuries previously described share common pathomechanics and clinical impairments (Table 32.1). Owing to this, treatment strategies are often similar across the different conditions. In addition, for

TABLE 32.1 Summary of chronic injuries. Injury mechanism pathway

Biomechanical risk factors

Impairments

Treatment

Chronic ankle instability

Macrotrauma with insufficient rehabilitation and/or presence of risk factors

Poor postural control; laterally displaced center of pressure during functional tasks [82]; reduced dorsiflexion range of motion

Impaired static and dynamic balance [30]; reduced ankle [83], knee, hip strength [84]; restricted dorsiflexion range of motion [85]

Balance training [86]; strength training [87]; restore biomechanics; restore range of motion [37]

Plantar fasciitis

Microtrauma with presence of risk factors or macrotrauma

Pes planus [88]; pes cavus; hindfoot varus; forefoot varus; reduced dorsiflexion range of motion [45]

Pain and discomfort with weight bearing [89]; restricted dorsiflexion [90]; restricted subtalar and midfoot range of motion [91]

Pain management [92]; rest with partial non weight-bearing; restore range of motion [93]; heel cups or orthotics [94]

Tendinopathy

Microtrauma with presence of risk factors or macrotrauma

Pes planus; pes cavus; forefoot varus; hindfoot varus/valgus; reduced dorsiflexion range of motion [95]

Pain with active and passive range of motion; reduced strength of involved musculature; altered biomechanics [96]

Eccentric strength training [97]; restore range of motion [98]; restore biomechanics

Stress fractures

Microtrauma with presence of risk factors

Altered muscle function [99]; anatomical abnormalities of metatarsals [100]; restricted subtalar motion

Pain and discomfort with weight-bearing; inflammation

Rest with nonweight bearing [64]; strengthening with appropriate tensile loading [65,101] (Continued )

TABLE 32.1 (Continued) Injury mechanism pathway

Biomechanical risk factors

Impairments

Treatment

Sesamoiditis

Microtrauma with presence of risk factors

Pes cavus; flexed first ray [75]; reduced dorsiflexion range of motion [76]

Pain with active and passive range of motion [75]; reduced intrinsic and extrinsic foot muscle strength

Restore range of motion; intrinsic foot strengthening; restore biomechanics

Retrocalcaneal bursitis

Microtrauma with presence of risk factors or macrotrauma

Bony abnormalities; flexed first ray; pes cavus; hindfoot varus; forefoot valgus

Pain; inflammation; reduced range of motion [102]

Pain and inflammation control [103]; restore biomechanics and structural malalignments [104]

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as much that risk factors were discussed within the context of injury, we must acknowledge that chronic injuries may or may not occur in the presence of risk factors. Finally, the focus of this chapter primarily presented impairments associated with the foot and ankle joint complex and disregarded proximal impairments and other risk factors (e.g., sex, body mass index, etc.). Emerging literature has found numerous proximal impairments, specifically altered hip strength and hip kinematics, which correspond with many of the chronic conditions presented.

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Arch Phys Med Rehab 2010;91(8):1267 71. [27] Hertel J. Functional anatomy, pathomechanics, and pathophysiology of lateral ankle instability. J Athl Train 2002;37(4):364 75. [28] Munn J, et al. Eccentric muscle strength in functional ankle instability. Med Sci sports Exerc 2003;35(2):245 50. [29] Arnold BL, et al. Ankle instability is associated with balance impairments: a meta-analysis. Med Sci Sports Exerc 2009;41(5):1048 62. [30] McKeon PO, Hertel J. Systematic review of postural control and lateral ankle instability, part I: can deficits be detected with instrumented testing? J Athl Train 2008;43(3):293 304. [31] Brown C. Foot clearance in walking and running in individuals with ankle instability. Am J Sports Med 2011;39(8):1769 76. [32] Chinn L, Dicharry J, Hertel J. Ankle kinematics of individuals with chronic ankle instability while walking and jogging on a treadmill in shoes. Phys Ther Sport 2013;14(4):232 9. [33] Doherty C, et al. Treatment and prevention of acute and recurrent ankle sprain: an overview of systematic reviews with meta-analysis. Br J Sports Med 2017;51(2):113 25. [34] Kaminski TW, et al. National Athletic Trainers’ Association position statement: conservative management and prevention of ankle sprains in athletes. J Athl Train 2013;48(4):528 45.

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[35] Kerkhoffs GM, et al. Diagnosis, treatment and prevention of ankle sprains: an evidence-based clinical guideline. Br J Sports Med 2012;46 (12):854 60. [36] McKeon PO, Donovan L. A perceptual framework for conservative treatment and rehabilitation of ankle sprains: an evidence-based paradigm shift. J Athl Train 2019;54(6):628 38. [37] Hoch MC, et al. Two-week joint mobilization intervention improves self-reported function, range of motion, and dynamic balance in those with chronic ankle instability. J Orthop Res 2012;30(11):1798 804. [38] McKeon PO, Hertel J. Systematic review of postural control and lateral ankle instability, part II: is balance training clinically effective? J Athl Train 2008;43(3):305 15. [39] Kosik KB, et al. Therapeutic interventions for improving self-reported function in patients with chronic ankle instability: a systematic review. Br J Sports Med 2017;51(2):105 12. [40] Lemont H, Ammirati KM, Usen N. Plantar fasciitis: a degenerative process (fasciosis) without inflammation. J Am Podiatr Med Assoc 2003;93 (3):234 7. [41] Roxas M. Plantar fasciitis: diagnosis and therapeutic considerations. Altern Med Rev 2005;10(2):83 93. [42] Wearing SC, et al. The pathomechanics of plantar fasciitis. Sports Med 2006;36(7):585 611. [43] Bolgla LA, Malone TR. Plantar fasciitis and the windlass mechanism: a biomechanical link to clinical practice. J Athl Train 2004;39(1):77 82. [44] Rome K, Howe T, Haslock I. Risk factors associated with the development of plantar heel pain in athletes. Foot 2001;11(3):119 25. [45] Riddle DL, et al. Risk factors for plantar fasciitis: a matched case-control study. J Bone Jt Surg Am 2003;85(5):872 7. [46] Pohl MB, Hamill J, Davis IS. Biomechanical and anatomic factors associated with a history of plantar fasciitis in female runners. Clin J Sport Med 2009;19(5):372 6. [47] Maganaris CN, et al. Biomechanics and pathophysiology of overuse tendon injuries: ideas on insertional tendinopathy. Sports Med 2004;34 (14):1005 17. [48] Bass E. Tendinopathy: why the difference between tendinitis and tendinosis matters. Int J Massage Bodyw 2012;5(1):14 17. [49] Maffulli N, Khan KM, Puddu G. Overuse tendon conditions: time to change a confusing terminology. Arthroscopy 1998;14(8):840 3. [50] Maffulli N, Wong J, Almekinders LC. Types and epidemiology of tendinopathy. Clin Sports Med 2003;22(4):675 92. [51] McShane JM, Ostick B, McCabe F. Noninsertional Achilles tendinopathy: pathology and management. Curr Sports Med Rep 2007;6 (5):288 92. [52] Maffulli N, Sharma P, Luscombe KL. Achilles tendinopathy: aetiology and management. J R Soc Med 2004;97(10):472 6. [53] Sayana MK, Maffulli N. Insertional Achilles tendinopathy. Foot Ankle Clin 2005;10(2):309 20. [54] Neville C, et al. Comparison of changes in posterior tibialis muscle length between subjects with posterior tibial tendon dysfunction and healthy controls during walking. J Orthop Sports Phys Ther 2007;37(11):661 9. [55] Wilder RP, Sethi S. Overuse injuries: tendinopathies, stress fractures, compartment syndrome, and shin splints. ClSports Med 2004;23 (1):55 81. [56] Woods L, Leach RE. Posterior tibial tendon rupture in athletic people. Am J Sports Med 1991;19(5):495 8. [57] Mosier SM, Pomeroy G, Manoli 2nd A. Pathoanatomy and etiology of posterior tibial tendon dysfunction. Clin Orthop Relat Res 1999;365:12 22. [58] Jackson MA, Gudas CJ. Peroneus longus tendinitis: a possible biomechanical etiology. J Foot Surg 1982;21(4):344 8. [59] Heckman DS, Gluck GS, Parekh SG. Tendon disorders of the foot and ankle, part 1: peroneal tendon disorders. Am J Sports Med 2009;37 (3):614 25. [60] Kulig K, et al. Effect of eccentric exercise program for early tibialis posterior tendinopathy. Foot Ankle Int 2009;30(9):877 85. [61] Rees JD, et al. The mechanism for efficacy of eccentric loading in Achilles tendon injury; an in vivo study in humans. Rheumatol 2008;47 (10):1493 7. [62] Silbernagel KG, et al. Full symptomatic recovery does not ensure full recovery of muscle-tendon function in patients with Achilles tendinopathy. Br J Sports Med 2007;41(4):276 80. [63] Mandell JC, Khurana B, Smith SE. Stress fractures of the foot and ankle, part 1: biomechanics of bone and principles of imaging and treatment. Skelet Radiol 2017;46(8):1021 9. [64] Ramponi DR, Hedderick V, Maloney SC. Metatarsal stress fractures. Adv Emerg Nurs J 2017;39(3):168 75. [65] Miller TL, Best TM. Taking a holistic approach to managing difficult stress fractures. J Orthop Surg Res 2016;11(1):98. [66] Lee S, Anderson RB. Stress fractures of the tarsal navicular. Foot Ankle Clin 2004;9(1):85 104. [67] Pavlov H, Torg JS, Freiberger RH. Tarsal navicular stress fractures: radiographic evaluation. Radiology 1983;148(3):641 5. [68] Hossain M, et al. Stress fractures of the foot. Clin Sports Med 2015;34(4):769 90. [69] Mandell JC, Khurana B, Smith SE. Stress fractures of the foot and ankle, part 2: site-specific etiology, imaging, and treatment, and differential diagnosis. Skelet Radiol 2017;46(9):1165 86. [70] Chuckpaiwong B, et al. Second metatarsal stress fracture in sport: comparative risk factors between proximal and non-proximal locations. Br J Sports Med 2007;41(8):510. [71] Hughes LY. Biomechanical analysis of the foot and ankle for predisposition to developing stress fractures. J Orthop Sports Phys Ther 1985;7 (3):96 101. [72] Fitch K, Blackwell J, Gilmour W. Operation for non-union of stress fracture of the tarsal navicular. J Bone Joint Surg Br Vol 1989;71 (1):105 10.

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[73] Boukhemis K, et al. Second metatarsal stress fractures. Oper Tech Orthop 2018;28(2):84 90. [74] Srinivasan R. The hallucal-sesamoid complex: normal anatomy, imaging, and pathology. Semin Musculoskelet Radiol 2016;20(2):224 32. [75] Boike A, Schnirring-Judge M, McMillin S. Sesamoid disorders of the first metatarsophalangeal joint. Clin Podiatr Med Surg 2011;28 (2):269 85. [76] Richardson EG. Injuries to the hallucal sesamoids in the athlete. Foot Ankle 1987;7(4):229 44. [77] Pekala PA, et al. The Achilles tendon and the retrocalcaneal bursa: an anatomical and radiological study. Bone Jt Res 2017;6(7):446 51. [78] van Dijk CN, et al. Terminology for Achilles tendon related disorders. Knee Surg Sports Traumatol Arthrosc 2011;19(5):835 41. [79] Canoso JJ, et al. Physiology of the retrocalcaneal bursa. Ann Rheum Dis 1988;47(11):910 12. [80] Sella EJ, Caminear DS, McLarney EA. Haglund’s syndrome. J Foot Ankle Surg 1998;37(2):110 14. [81] Stephens MM. Haglund’s deformity and retrocalcaneal bursitis. Orthop Clin North Am 1994;25(1):41 6. [82] Schmidt H, et al. Increased in-shoe lateral plantar pressures with chronic ankle instability. Foot Ankle Int 2011;32(11):1075 80. [83] Arnold BL, et al. Concentric evertor strength differences and functional ankle instability: a meta-analysis. J Athl Train 2009;44(6):653 62. [84] Friel K, et al. Ipsilateral hip abductor weakness after inversion ankle sprain. J Athl Train 2006;41(1):74 8. [85] Denegar CR, Hertel J, Fonseca J. The effect of lateral ankle sprain on dorsiflexion range of motion, posterior talar glide, and joint laxity. J Orthop Sports Phys Ther 2002;32(4):166 73. [86] McKeon PO, et al. Balance training improves function and postural control in those with chronic ankle instability. Med Sci Sports Exerc 2008;40(10):1810 19. [87] Docherty CL, Moore JH, Arnold BL. Effects of strength training on strength development and joint position sense in functionally unstable ankles. J Athl Train 1998;33(4):310 14. [88] Prichasuk S, Subhadrabandhu T. The relationship of pes planus and calcaneal spur to plantar heel pain. Clin Orthop Relat Res 1994;306:192 6. [89] Phillips A, McClinton S. Gait deviations associated with plantar heel pain: a systematic review. Clin Biomech (Bristol, Avon) 2017;42:55 64. [90] McPoil TG, et al. Heel pain plantar fasciitis: clinical practice guildelines linked to the international classification of function, disability, and health from the orthopaedic section of the American Physical Therapy Association. J Orthop Sports Phys Ther 2008;38(4):A1 18. [91] Irving DB, Cook JL, Menz HB. Factors associated with chronic plantar heel pain: a systematic review. J Sci Med Sport 2006;9(1 2):11 22. [92] Diaz Lopez AM, Guzman Carrasco P. Effectiveness of different physical therapy in conservative treatment of plantar fasciitis: systematic review. Rev Esp Salud Publica 2014;88(1):157 78. [93] Young CC, Rutherford DS, Niedfeldt MW. Treatment of plantar fasciitis. Am Fam Physician 2001;63(3):467 74 477-8. [94] Uden H, Boesch E, Kumar S. Plantar fasciitis - to jab or to support? A systematic review of the current best evidence. J Multidiscip Healthc 2011;4:155 64. [95] Rabin A, Kozol Z, Finestone AS. Limited ankle dorsiflexion increases the risk for mid-portion Achilles tendinopathy in infantry recruits: a prospective cohort study. J Foot Ankle Res 2014;7(1):48. [96] Bubra PS, et al. Posterior tibial tendon dysfunction: an overlooked cause of foot deformity. J Family Med Prim Care 2015;4(1):26 9. [97] Murtaugh B, Ihm JM. Eccentric training for the treatment of tendinopathies. Curr Sports Med Rep 2013;12(3):175 82. [98] Duthon VB, et al. Noninsertional Achilles tendinopathy treated with gastrocnemius lengthening. Foot Ankle Int 2011;32(4):375 9. [99] Bennell K, et al. Risk factors for stress fractures. Sports Med 1999;28(2):91 122. [100] Korpelainen R, et al. Risk factors for recurrent stress fractures in athletes. Am J Sports Med 2001;29(3):304 10. [101] Gehrmann RM, Renard RL. Current concepts review: stress fractures of the foot. Foot Ankle Int 2006;27(9):750 7. [102] Chimenti RL, et al. Patients with insertional Achilles tendinopathy exhibit differences in ankle biomechanics as opposed to strength and range of motion. J Orthop Sports Phys Ther 2016;46(12):1051 60. [103] Chimenti RL, et al. Current concepts review update: insertional Achilles tendinopathy. Foot Ankle Int 2017;38(10):1160 9. [104] Wiegerinck JI, et al. Treatment for insertional Achilles tendinopathy: a systematic review. Knee Surg Sports Traumatol Arthrosc 2013;21 (6):1345 55.

Chapter 33

Hallux Valgus Sheree Hurn Faculty of Health, School of Clinical Sciences, Podiatry, Queensland University of Technology (QUT), Brisbane, QLD, Australia

Abstract Hallux valgus (HV) is a prevalent musculoskeletal foot deformity. The cause of HV is uncertain and multifactorial, with proposed factors including genetics, footwear, foot structure, and biomechanics. HV may be diagnosed clinically and assessed using radiography, ultrasound, or three-dimensional imaging methods. Patients present with a range of concerns including foot pain, difficulty fitting footwear, and concerns about appearance. Functional deficits in moderate-to-severe cases may include weakness and atrophy of the toe flexor muscles, increased postural sway, altered plantar pressures, and in older adults, gait instability, and increased risk of falls. A range of nonsurgical treatment options may be trialed prior to surgical correction of HV, although there is limited evidence for their effectiveness. These options include foot orthoses, night splints, toe separators, manual therapies, and foot exercises. Longterm prospective studies are needed to investigate risk factors for HV and the efficacy of nonsurgical treatments.

33.1

Introduction

Hallux valgus (HV) is a common musculoskeletal foot deformity, presenting with lateral deviation of the hallux toward the lesser toes, medial deviation of the first metatarsal head, and subluxation of the first metatarsophalangeal (MTP) joint. A painful soft tissue bursa may develop over the medial prominence of the first metatarsal head, commonly referred to as a “bunion.” Despite this distinction, the terms HV and bunion are sometimes used interchangeably, and the term “bunion” is more often used in lay terms when referring to HV. HV affects all age groups, and when observed prior to skeletal maturity, it may be referred to as juvenile HV, pediatric HV, or adolescent bunion [1]. This chapter outlines the following: prevalence and etiology of HV; diagnosis and imaging for HV assessment; clinical presentation in terms of pain, footwear and function; functional outcomes, including balance and gait analysis studies; and finally, nonsurgical and, briefly, surgical treatment pathways.

33.2

Prevalence

HV is very common in the general population, with a meta-analysis by Nix et al. [2] reporting a prevalence of 23% in adults aged 18 to 65 years. HV is increasingly common with age, affecting 36% of older adults aged over 65 years. This meta-analysis also found a significantly higher prevalence of HV in women (30%) compared to men (13%) [2], and this is often observed clinically. Although HV is highly prevalent, it is clear from the literature that only a proportion of these individuals progress to having severe deformity or experience foot pain requiring treatment. Out of 60 adults with HV recruited from the community, 16 (27%) were classed as having disabling foot pain by Hurn et al. [3]. Cho et al. [4] reported that only 43 out of 364 participants with HV (12%) had foot pain. Painful HV is more likely in populations seeking treatment, with 39 out of 94 patients presenting for HV surgery (41%), experiencing associated metatarsalgia [5]. Studies have used varying definitions of foot pain and severity of deformity, which makes it difficult to compare estimates. Cho et al. [4] reported that out of 364 participants with HV, 48 (13%) were classed as having a moderate or greater degree of deformity (.25 degrees), whereas in the study by Hurn et al. [3], 26 participants out of 60 (43%) were classed as having moderate deformity and 24 participants (40%) had severe deformity based on the Manchester Scale. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00015-9 © 2023 Elsevier Inc. All rights reserved.

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BOX 33.1 Proposed risk factors for hallux valgus Systemic factors G Genetics and race [8 11] G Age and sex (HV is more common in women and older adults) [2] G Systemic disease [6,7] Structural and biomechanical factors G Long first metatarsal [12] G Round first metatarsal head [12] G Increased first intermetatarsal (IM) angle [12] G Generalized ligamentous laxity [13] G First ray hypermobility [14,15] G Pes planus [16 22] G Tight Achilles tendon [23 25] G Intrinsic muscle imbalance around the first MTP joint [26 32] Extrinsic factors G Footwear (tight, pointed toe box, or high heeled) [33 39]

33.3

Etiology

The etiology and pathophysiology of HV are complex and multifactorial. A wide range of potential underlying causes for HV have been proposed, including genetics, biomechanical factors, and shoe-wearing habits. In some cases, the joint subluxation associated with HV may be secondary to a systemic condition such as rheumatoid arthritis [6] or a neuromuscular disorder [7]. The majority of studies to date investigating HV etiology are cross-sectional, with no high-quality prospective studies, therefore conclusions regarding causation or risk factors for HV cannot be drawn. Nonetheless, a number of cross-sectional studies have reported statistically significant associations between various factors and HV, and expert opinion papers have proposed potential theoretical mechanisms on how these factors may lead to HV development, based on biomechanical principles. The following paragraphs will present the most commonly discussed theoretical mechanisms underlying HV development and their significant associated factors (see Box 33.1).

33.3.1 Genetics and race It is often anecdotally reported by patients and clinicians that HV is an inherited condition. Studies have estimated that a family history of HV is present in 63 to 90% of cases [9,10]. Hannan et al. [8] reported significant heritability of HV, based on data from the Framingham Foot Study, which included an adult population (n 5 1370) of European descent. In this study sample, heritability for HV was moderate to high, depending on age and sex. Another study has shown racial differences in the occurrence of HV, with African Americans more likely than white Americans to have HV in a sample of 1691 participants [11].

33.3.2 Structural and biomechanical factors Several structural features of the foot have been reported to be associated with HV. A systematic review [12] has reported statistically significant associations between HV and a longer first metatarsal, a round first metatarsal head, increased first IM angle, and lateral sesamoid displacement. Although these cross-sectional studies cannot determine causality, it could be hypothesized that a longer first metatarsal with a round-shaped metatarsal head could be inherited features that may predispose an individual to developing HV, whilst an increased IM angle and lateral displacement of the sesamoids may be secondary features of HV deformity. Generalized ligamentous laxity [13] and first ray hypermobility have both been discussed in relation to HV development [14,15]. Perera et al. [25] outline first ray biomechanics in HV, highlighting that as there are no tendon attachments on the head of the first metatarsal, stability and congruence of the first MTP joint and tarsometatarsal joint during gait are vital. These authors propose that failure at any point along the first ray may result in the development of HV.

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Another biomechanical factor often discussed in relation to HV development is pes planus or a lowered arch profile. This mechanism was proposed by Inman in 1974 [19], although not substantiated by quantitative data but rather by anecdotal observations. Inman noted that hindfoot eversion in weight-bearing coincides with longitudinal rotation of the first ray, placing the first ray axis in an oblique and unstable position, “less able to withstand deforming pressures exerted upon it.” Studies investigating the link between HV and static measures of arch height (calcaneal angle, navicular height, arch index) have not shown any statistically significant associations [16,17,20 22]. However, considering dynamic function during gait may provide more insights in relation to HV development (see Gait analysis section described later in this chapter). Data from the Framingham Foot Study [18] showed increased odds of having HV in participants with a lower center of pressure excursion index (indicative of feet functioning with more hindfoot eversion and medial longitudinal arch collapse). Interestingly, the same study showed no significant association between HV and a pes planus foot type, measured in static stance using the modified arch index [18]. A tight Achilles tendon is another biomechanical factor proposed to be associated with HV [24,25], although there is limited and inconsistent evidence to support this association [24]. The theoretical link is based on increased and early forefoot loading during the gait cycle, which may lead to external rotation of the foot and increased force on the medial border of the hallux [25]. Clinical observations suggest that gastrocnemius tightness may be of particular importance in the development of juvenile HV. In a case series reported by Barouk [23], 77% of patients presenting with gastrocnemius tightness had concomitant HV, and 71% of these had experienced a juvenile pattern of HV development. Intrinsic muscle imbalance between the abductor and adductor muscles surrounding the first MTP joint is proposed as another factor in HV development. Conversely, some morphological changes and muscle weakness may occur secondarily to HV deformity. Altered morphology of the abductor hallucis muscle in HV has been demonstrated via histological [27] and ultrasound investigations [30,32], and the abductor hallucis insertion has been shown to be in a more plantar position in those with HV deformity [26]. Studies using electromyography (EMG) have also demonstrated weakness and imbalance of the intrinsic foot muscles in HV participants compared to controls [27 29,31]. Extrinsic muscles crossing the first MTP joint also have a role to play, with the “bowstring” effect of the long flexor and extensor tendons clearly demonstrated on magnetic resonance imaging (MRI) [26].

33.3.3 Footwear HV is often attributed to shoe wearing habits of patients in earlier life, especially ladies fashion footwear with a high heel and narrow toe box; however, there are few high quality studies to support this hypothesis [35]. A study by Sim-Fook and Hodgson [39], which reported a higher prevalence of HV (33%) in a shoe wearing population (n 5 118) compared to a barefooted population (2%, n 5 107), is often cited in support of this theory; however, there were other significant differences between the two populations, which may have confounded these findings (e.g., age and gender). A more recent study by Choi et al. [34] showed that a group of partially shod Maasai women (n 5 20) had significantly lower HV angles compared to a group of regularly shod Maasai women (n 5 20), although the reported mean HV angles were nonpathological (,15 degrees). Klein et al. [40] reported a significant association between inadequate footwear length in preschool children and hallux angle (n 5 858), thus highlighting the importance of correct shoe fit in young children. Other studies have relied upon participants’ self-report data regarding past shoe wearing habits. While the impact of recall bias must be considered, two studies have shown HV to be more prevalent in women who report wearing shoes with a narrow toe box between the ages of 20 and 39 [33,41]. There is conflicting evidence regarding the potential association between HV and past history of wearing high heeled footwear [36,37]. Further studies are needed with validated and reliable methods to investigate these associations.

33.4

Diagnosis and imaging

33.4.1 Clinical diagnosis HV may be diagnosed clinically using noninvasive methods. The Manchester Scale is a validated classification scale used to grade HV as mild, moderate, or severe based on standardized photographs [42] (Fig. 33.1). Clinical photography may also be used to quantify HV angle [43] (Fig. 33.2). Grading the severity of deformity is important for patient education and also for monitoring progression of HV. Mild HV is very common and in asymptomatic cases may have minimal impact on foot function; whereas, moderate to severe HV is associated with impaired function. Functional outcomes in HV will be discussed in more detail later in this chapter.

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FIGURE 33.1 Manchester Scale for the grading of hallux valgus. (A) None. (B) Mild. (C) Moderate. (D) Severe. Reproduced with permission from Garrow et al. The grading of hallux valgus. The Manchester Scale. J Am Podiatr Med Assoc 2001;91(2):77.

Relevant medical history should be elicited from the patient, as HV may be associated with systemic conditions such as rheumatoid arthritis [6]. Furthermore, a simple measure of first MTP joint dorsiflexion range of motion should be obtained using a tractograph or goniometer [44], to differentiate HV from other conditions affecting the first MTP joint such as hallux rigidus or hallux limitus, which will present with reduced range of motion at the first MTP joint.

33.4.2 Radiographic assessment Plain radiographs are traditionally used to assess HV deformity, including weight-bearing dorsoplantar, lateral, and oblique views. Dorsoplantar views are often used for measuring weight-bearing alignment, with the most common angles including HV angle, IM angle, interphalangeal angle, and distal metatarsal articular angle [45] (Table 33.1 and Fig. 33.3). Radiographs for measurement purposes should always be taken using standardized methods according to published recommendations to improve reliability, with the X-ray tube angled 15 degrees from vertical, and the X-ray beam centered on the midfoot [45]. With respect to obtaining measurements from radiographs, five different methods have been described for bisecting the first metatarsal shaft; however, the consensus of the ad hoc committee of the

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FIGURE 33.2 Clinical photograph with measurement of hallux valgus angle.

TABLE 33.1 Angular measurements commonly taken from weight-bearing radiographs in hallux valgus [47]. Radiographic angle

Abbreviation

Normal range (degrees)

Definition

Hallux valgus angle

HVA

,15

Angle between axes of first metatarsal and hallux proximal phalanx. Reference points placed on proximal and distal midmetaphyseal regions.

1 2 Intermetatarsal angle

IMA

,9

Angle between axes of first and second metatarsals. Reference points on proximal and distal midmetaphyseal regions.

Distal metatarsal articular angle

DMAA

,7

Angle between first metatarsal axis and another line drawn as follows: reference points placed on most medial and lateral aspects of first metatarsal distal articular surface; these points connected with a straight line and another line drawn perpendicular.

Interphalangeal angle

IPA

Angle between axes of proximal and distal phalanges of hallux. Reference points on proximal phalanx as for HVA. Reference points for distal phalanx are distal tip and midpoint of the articular surface of the distal phalanx.

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FIGURE 33.3 Angular measurements commonly taken from weight-bearing radiographs in hallux valgus. HVA, hallux valgus angle; IMA, intermetatarsal angle; DMAA, distal metatarsal articular angle; IPA, interphalangeal angle.

American Orthopedic Foot and Ankle Society [46] proposed repeatable reference points on the first metatarsal diaphysis (Fig. 33.3). Radiographic classification of mild, moderate, and severe HV is based on both HV and IM angles, although there is a lack of consistency in the literature surrounding these classification scales. Hecht and Lin [48] propose the following classification based on standing dorsoplantar radiographs: G G G

Mild: HV angle 15 19 degrees and IM angle 9 11 degrees, ,50% subluxation of lateral sesamoid Moderate: HV angle 20 40 degrees, IM angle ,16 degrees, 50% 70% subluxation of lateral sesamoid Severe: HV angle .40 degrees, IM angle $ 16 degrees, .75% subluxation of lateral sesamoid

Further assessment of radiographs should include observation for signs of degenerative changes in the joint such as osteophytes and joint space narrowing. A radiographic atlas has been developed by Menz et al. [49], which has demonstrated moderate to excellent intra-rater reliability in identifying radiographic features of osteoarthritis in commonly affected joints of the foot. These authors found that using a dorsoplantar view alone detected 95% of cases of first MTP joint osteoarthritis.

33.4.3 Ultrasound Diagnostic ultrasound may be a useful tool for quantifying morphologic changes to the forefoot in all three body planes, in terms of bone and joint structures as well as soft tissue changes. Matsubara et al. [50] showed ultrasound to be reliable in measuring several structural parameters of the forefoot in the coronal plane: height of the second metatarsal,

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FIGURE 33.4 Ultrasound images of (A) cross-sectional area of abductor hallucis (left), and (B) dorsoplantar thickness of abductor hallucis muscle in an individual with moderate hallux valgus (right).

transverse arch height, angle of sesamoid rotation, and distance between medial sesamoid and fifth metatarsal head. Their study also showed good agreement between ultrasound measures and computed tomography (CT). Furthermore, HV is associated with changes involving the intrinsic and extrinsic muscles inserting around the hallux. Since musculature cannot be examined on plain radiographs, diagnostic ultrasound techniques may be helpful to visualize and quantify these changes. The abductor hallucis and flexor hallucis brevis muscles have been shown to have significantly reduced thickness and cross-sectional area in individuals with HV compared to those without HV [30,32,51]. Ultrasound methods have been shown to be reliable for this purpose (ICCs 0.79 to 0.98) [52] and thus may be a useful quantitative measure to evaluate the effectiveness of interventions that target muscle strengthening for individuals with HV [53]. Ultrasound imaging of an individual with moderate HV, demonstrating cross-sectional area and dorsoplantar thickness of the abductor hallucis muscle is presented (Fig. 33.4).

33.4.4 Computed tomography Traditional two-dimensional imaging methods have limitations in their ability to represent the complex threedimensional alignment of the first ray segments in HV. Therefore, three-dimensional imaging techniques may provide further insights into the pathomechanics of HV. Whilst conventional cross-sectional imaging techniques are not weightbearing, some researchers have utilized a device to load the foot during CT examination, and Welck et al. [54] used three-dimensional, weight-bearing cone beam CT to investigate HV. Collan et al. [55] first reported weight-bearing CT data for a group of HV patients compared to controls and these authors reported high correlation with traditional radiographic measures. Two studies have used CT and three-dimensional modeling to demonstrate first ray hypermobility and alignment in HV patients compared to controls [56,57], demonstrating adduction, dorsiflexion, and inversion of the first metatarsal with respect to the medial cuneiform in HV. Weight-bearing CT has also been demonstrated to be highly reliable for evaluating the sesamoid complex in HV, including sesamoid position, rotation, and metatarsosesamoid joint space [54].

33.4.5 Magnetic resonance imaging MRI may reveal further insights into the soft tissue and tendon structures surrounding the first ray in HV, while at the same time demonstrating alignment of the first ray in three body planes. Schweitzer et al. [58] observed that medial eminence thickening (95%) and adventitious bursa (70%) around the first metatarsal head were common MRI findings in HV, as well as lateral sesamoid position (83%) and sesamoid proliferation (90%). These authors also used MRI to observe several other features of HV commonly assessed on plain radiographs, for example metatarsal head shape, first metatarsal length, signs of first MTP joint osteoarthritis, and flattening of the longitudinal arch. It has been reported

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from an MRI study that the abductor hallucis tendon is displaced in a plantar direction in patients with HV [26], and this tendon shift is correlated with the severity of HV angle. Furthermore, the flexor hallucis longus tendon is laterally displaced, and this can also be demonstrated on MRI [59]. MRI studies can demonstrate cross-sectional area of the intrinsic foot muscles in a similar way to ultrasound imaging, and this parameter on MRI has been correlated with toe flexion strength [60]. Weight-bearing MRI has been utilized by Glasoe et al. [61,62] in simulated positions representing key point in the gait cycle, and their findings will be discussed later in this chapter.

33.5

Clinical presentation

33.5.1 Foot pain Clinical assessment of HV must include a thorough history of the patient’s presenting concerns, as this may guide treatment decisions. HV deformity may present as a relatively pain-free condition, with the patient’s primary concern being about the appearance of their feet. This is common in juvenile HV, alongside concerns about family history and progression of the deformity [63]. Foot pain is reported to be a very common presenting concern associated with HV [63]. However, there is no correlation between increasing lateral deviation of the hallux and severity of self-reported foot pain in HV [3]. There are conflicting reports from community-based studies regarding whether or not the presence of HV is associated with foot pain. A number of studies have reported no association between HV and foot pain [37,64 67], while other community studies have linked the presence of HV to foot pain [4,68 71]. Different methods for determining the presence and severity of foot pain and HV can help explain these inconsistent reports. Of these community studies, those using validated tools for HV diagnosis [70,71] and validated foot-specific pain scales [4,70] have demonstrated positive associations between HV and foot pain. Where pain is present in HV, it may be associated with various anatomical structures of the forefoot, and the clinician must determine from a thorough history and examination which structures are likely to be involved. Pain associated with HV may be due to nerve impingement of the dorsal or plantar digital nerves as they pass over the medial exostosis, resulting in a description of sharp, burning pain or paresthesia [72]. Another distinct type of pain accompanying HV may be bursal inflammation, associated with shoe pressure and noticeable redness over the medial exostosis [72]. Pain may be due to osteoarthritic changes in the first MTP joint or sesamoid complex [17], and this pain may be reproduced with movement and/or palpation of the first MTP joint and sesamoids. Finally, pain may be associated with secondary lesions or hyperkeratosis, and this is especially common in older adults [38].

33.5.2 Footwear Another common presenting concern in patients with HV is regarding difficulty fitting footwear [38,63]. A key part of the initial assessment should include a thorough assessment of the patient’s footwear. Features that have been discussed in the literature as potentially associated with HV include: a heel height greater than 25 mm, narrow pointed toe box, or simply inadequate width [12]. Expert consensus is to avoid tight-fitting and high heeled footwear which may exacerbate symptoms [48], thus shoe features and fit must both be considered. Footwear advice is likely to form the basis of any conservative treatment plan for HV [72].

33.5.3 Self-reported function and quality of life Functional limitations may be a presenting concern associated with HV [63]. HV is associated with poorer self-reported function [4,38,70] and quality of life [73]. Menz et al. [70] demonstrated that increasing severity of HV has a greater impact on general and foot-specific health-related quality of life, controlling for age, sex, education, and BMI. Even though there was no difference between HV and control participants in terms of general health (determined using the SF-36v2 health survey), HV participants reported significantly poorer foot function [38].

33.6

Functional outcomes

33.6.1 Balance and falls HV has been linked to an increased risk of falls in older adults [74 78] and poorer performance in functional tests such as the coordinated stability and lateral stability tests [79,80]. Hurn et al. [81] showed increased mediolateral postural

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sway in single leg stance in adults (mean age 49 years) with severe HV compared to controls. Menz et al. [82] showed a significant correlation between HV severity score and poorer performance in several balance tasks in older adults. Based on these findings, it appears that significantly impaired balance may only be observed in cases of severe HV. Furthermore, it should be noted that there are conflicting reports in the literature, with Mickle et al. [83] finding no significant difference in postural sway between older adults with HV and controls. One factor that may help explain potential balance impairment associated with HV is the size of the abductor hallucis muscle, which is known to be reduced in HV. Zhang et al. [84] reported an inverse correlation between the size of the abductor hallucis muscle and center of pressure sway parameters, although their study was not conducted in a population with HV. Falls risk assessment and balance exercises should be recommended in older adults with HV in view of the increased falls risk in this population.

33.6.2 Hallux flexion and abduction Hallux flexion strength is an important clinical measure due to its association with falls risk in older adults [76,77] and the necessity of toe flexion for maintaining balance during forward reaching tasks [85]. Two studies have shown significantly reduced hallux flexion force, measured using load cell and pressure platform methods, in those with moderate to severe HV compared to those without HV [77,81] Another study [86] has shown an inverse correlation between hallux flexion strength and HV angle. Toe flexor strength has been shown to be correlated with cross-sectional area of the plantar intrinsic foot muscles in young healthy participants [60]. As discussed previously, the cross-sectional area of the abductor hallucis and flexor hallucis brevis are reduced in HV, and this could help explain the reduced toe flexion strength associated with HV. Studies using EMG methods have revealed another aspect of foot function in HV. Iida and Basmajian [28] found that during hallux flexion tasks (seated and standing), intrinsic muscle activity was similar in both HV and control groups. However, during tip-toe standing, the flexor hallucis brevis and abductor hallucis muscles were more active (expressed as a percentage of maximum voluntary contraction) in the HV group compared to controls. Three EMG studies have shown reduced abductor hallucis muscle activity during abduction tasks [28,29,87] but two of these studies demonstrated increased abductor hallucis activity during flexion [28,29], indicating that the abductor hallucis muscle becomes more of a flexor due to plantar displacement of the tendon that occurs in HV deformity [26]. There is debate surrounding whether muscle weakness develops secondarily to the deviation of the hallux, or whether muscle imbalance is a primary aspect of HV development. Hoffmeyer et al. [27] in a case series took muscle biopsies from HV patients undergoing surgery, identifying abnormalities in 53 out of 57 specimens. The authors hypothesized that this could be due to localized ischemia, associated with acute compartment syndromes developing within the firmly sheathed intrinsic foot muscles. Stewart et al. [32] noted reduced cross-sectional area of abductor hallucis and flexor hallucis brevis in mild HV, and no significant differences were found between mild, moderate, and severe HV. These authors concluded that changes in the abductor hallucis muscle are present early in the development of HV, and may therefore predispose an individual to progression of HV. However, Hurn et al. [81] found no significant difference in hallux flexion and abduction strength between those with mild HV and controls, indicating that muscle weakness may develop as a secondary feature of HV. The abductor hallucis muscle in particular has been shown to be very important for balance and postural tasks [88] and further research is warranted to understand the link between reduced hallux flexion strength, postural instability, and intrinsic foot muscle activity.

33.6.3 Gait analysis Studies have shown gait alterations in adults with HV, in terms of kinematics, plantar pressures, muscle activity, and spatiotemporal parameters [89].

33.6.3.1 Kinematics In terms of lower limb kinematics, one study using the Oxford foot model showed that during the terminal stance phase of gait, individuals with HV had reduced ankle dorsiflexion and reduced internal rotation of the hindfoot with respect to the tibia [90]. Another kinematic analysis of HV by Canseco et al. [91] using a four-segment foot model showed less flexion of the forefoot with respect to the hindfoot throughout the gait cycle, indicating flattening of the medial longitudinal arch. Hwang et al. [92] described a nine-segment foot model and although their study only included two participants with HV, compared to 10 individuals without HV, their data showed a trend toward greater eversion at the subtalar joint during early stance, and greater talocrural joint eversion during terminal stance. The findings of these studies are consistent with theoretical mechanisms of HV development described earlier (ankle equinus and pes planus).

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Glasoe et al. [62,93] utilized weight-bearing MRI studies to investigate the triplanar movement of the tarsal bones in simulated positions of gait (mid-stance, heel-off, and terminal stance) in participants with and without HV. They reported an incremental increase in both dorsiflexion and abduction at the first MTP joint between the three gait positions. These studies found a significant difference in first ray axis inclination between study participants with and without HV. This increase in inclination of the first ray axis results in lowering of the navicular and increased adduction of the first metatarsal. As advances in modern gait analysis and imaging techniques create more opportunities for threedimensional imaging of the first MTP joint both statically and dynamically, these techniques may be very informative to help increase understanding of first ray biomechanics in individuals with HV.

33.6.3.2 Plantar pressures There are inconsistent reports regarding plantar pressure changes in HV, with some studies showing increased plantar pressures under the hallux [94,95] and first-second metatarsals [83,94]; however, another study showed reduced plantar pressures under the first metatarsal [96]. A study by Hurn et al. [81] explored these changes in mild, moderate, and severe HV, reporting that hallux pressures were significantly reduced in moderate and severe HV groups. This is consistent with two other studies that have showed an inverse relationship between hallux pressures and HV angle [97,98]. Hurn et al. [81] showed that differences in first metatarsal peak pressures were not significant after adjusting for age, sex, body mass index, and foot pain. These authors propose that pain under the first MTP joint may be the reason that forefoot loading would be altered, rather than hallux deviation itself. This would explain the inconsistent findings from previous studies, as some study populations may have had painful HV while others may have been asymptomatic. It is also evident that severity of HV angle deviation should be considered, as significant gait changes may not be evident in mild HV.

33.6.3.3 Muscle activity There have been limited studies investigating EMG activity of muscles during gait in individuals with HV, although muscle imbalance is thought to be a significant factor in HV development. Shimazaki and Takebe [31] showed that individuals with HV were more likely to show simultaneous onset at heel strike of the intrinsic muscles examined using fine wire EMG (abductor hallucis, adductor hallucis, flexor hallucis brevis, and extensor hallucis brevis). Hoffmeyer et al. [27] observed that individuals with HV compared to controls were more likely to exhibit abnormal activation of the abductor hallucis and first dorsal interosseous muscles during gait. Abnormal activity patterns were considered to be either negligible activity, or continuous muscle activity throughout stance and swing, while bursts of muscle activity, considered to be the normal pattern, were more common in the control group.

33.6.3.4 Temporospatial parameters A number of studies have shown no difference in spatiotemporal parameters (gait cycle time, percent stance/swing, cadence, speed, and step length) between HV and control groups [83,90,95]. Taranto et al. [22] showed no difference in toe-out angle between HV and control groups. However, one study by Canseco [91] showed slower walking speed, decreased stride length, and prolonged stance phase in HV patients compared to controls. Another study by Menz and Lord [99] found that older adults with HV when walking on an irregular surface had reduced walking speed, and a shorter step length when walking on either a level or irregular surface compared to controls. The same study by Menz and Lord [99] found lower vertical harmonic ratios at the head and pelvis during walking on an irregular surface, which indicates a less stable gait pattern in older adults with HV compared to controls.

33.7

Treatment pathways

33.7.1 Nonsurgical treatment The wide range of nonsurgical treatment options for HV is outlined in Box 33.2 and below. Conservative treatments should be trialed by patients with HV before considering surgical options [48,72]. There will be patients who are not appropriate surgical candidates due to medical comorbidities, for whom conservative treatments will be the best option. Nonsurgical options are also recommended in cases of juvenile HV, due to the high risk of recurrence and possibility of overcorrection if surgery is performed before skeletal maturity [48]. The various goals of nonsurgical treatment often include pain reduction, prevention of deformity progression, accommodation of deformity, and offloading any secondary pressure lesions [100].

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BOX 33.2 Nonsurgical treatment options for hallux valgus Expert opinion recommendations often include the following: G Footwear advice and modification: wide forefoot, low heel height, extra depth accommodative upper made of flexible material [48,72,100,101] G Bunion shield padding [72] G Cryotherapy [72] G Antiinflammatory medications [72] G Foot orthoses or insoles: full length, medially posted, with metatarsal pad or bunion flare [72,100,101] G Night splinting to apply adductory stretch to the first MTP joint capsule [100] G Toe spacers [100,101] Other treatments options include physical therapies and pharmacological approaches: G Manual therapy [102,103] G Taping [104] G Foot exercise (e.g., toe spread) [53] G Botulinum toxin A injections [105]

FIGURE 33.5 Silicone bunion shield.

33.7.1.1 Expert opinion and current practice There is a lack of quality scientific evidence to support the effectiveness of nonsurgical interventions for HV [106]; however, expert opinion papers offer a consensus that footwear advice and modification often form the basis of nonsurgical treatment for HV [48,72,100,101]. Bunion shield padding (Fig. 33.5), cryotherapy and antiinflammatory medications may be offered for symptomatic relief [72]. Foot orthoses (Fig. 33.6), night splints (Fig. 33.7) and toe separators (Fig. 33.8) are also often recommended [72,100,101]. Original research articles investigating the effectiveness of these interventions are outlined in the following sections. A survey of current practice by Australian podiatrists treating HV found that a wide range of nonsurgical treatment options are utilized in practice [63]. In accordance with expert opinion, the most common treatments identified were footwear advice and modification, as well as foot orthoses. However, this survey identified significant differences between patient age groups in terms of presenting concerns and treatment recommendations for HV. In cases of juvenile HV, podiatrists were more likely to recommend prefabricated foot orthoses and exercises, whereas custom foot orthoses

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FIGURE 33.6 Foot orthoses for hallux valgus.

FIGURE 33.7 Hallux valgus night splint.

were more often prescribed in adults with HV, and bunion shields or shoe padding were more likely to be recommended in older adults with HV. In future, clinical guidelines for HV treatment should consider different treatment pathways for different age groups.

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FIGURE 33.8 Silicone toe spacer for hallux valgus.

33.7.1.2 Foot orthoses One randomized controlled trial conducted by Kilmartin et al. [107], including 122 children, investigated the effect of foot orthoses on HV angle. After three years of followup, the HV angle had progressed in both orthoses and control groups, indicating that the orthoses did not prevent the progression of HV in children. A small comparative study by Reina et al. [108] investigated the effect of custom foot orthoses compared to no treatment in 54 women, finding no significant differences in HV angle between baseline and followup and no significant differences between groups. Another randomized controlled trial by Torkki et al. [109] investigated the effects of foot orthoses in adults with painful HV (n 5 209). These authors reported a significant reduction in foot pain at six months in the orthoses group, compared to a control group of patients who were on the waiting list for surgery. However, the difference between groups was not significant at 12-months followup. Another study in patients with rheumatoid arthritis showed that those treated with custom foot orthoses were 73% less likely to develop HV during the three-year study followup period [110]. From the limited studies available, it appears that foot orthoses have no effect on HV angle, but may be useful for short-term relief of pain symptoms associated with HV, and may prevent the development of HV associated with rheumatoid arthritis.

33.7.1.3 Splints and toe separators Two case series reports have discussed the use of night splints in children [111,112]. Both reports had widely varied followup periods (range 18 months to 6 years), and no control or comparison group. Macfarlane et al. [112] reported no significant difference in radiographic HV angle between baseline and followup, concluding that night splint therapy may prevent the progression of HV in children. Groiso et al. [111] conducted no statistical analyses and therefore study findings are difficulty to interpret. Another study [113] compared a specially designed slipper with hallux splint to commercially available night splints and toe separators. The differences in HV angle reported over the 12-month study period were not clinically significant in either group (,0.3 ). Two small studies have reported the use of custom foot orthoses with a fixed toe separator. Tehraninasr et al. [114] reported significant pain reduction in the insole group compared to a group using night splints; however, there was no significant effect on HV angle in either group over a threemonth followup period. The study groups were randomized in their study but no blinding occurred. A case series report by Tang et al. [115] reported the use of a similar device, with significant improvement in pain and walking ability noted over three months. However, only 12 participants completed the study and there was no control or comparison group.

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BOX 33.3 Key points regarding nonsurgical treatment for hallux valgus G G G

No evidence that orthoses or night splint therapy have an effect on HV angle Orthoses and toe separators may improve pain symptoms in the short term More studies are needed investigating physical therapies such as manual therapy, taping, and exercise

33.7.1.4 Manual therapy Two studies have investigated the effects of manual and manipulative therapy techniques in adults with painful HV. One single-blind randomized controlled trial (n 5 60) reported an improvement in pain and functional outcomes in the short-term, following mobilization with cryotherapy, compared to a placebo [102]. Another study comparing manual therapy and night splints found significant short-term improvements in pain and function across both groups, although the small sample size (n 5 30) was under-powered to investigate between-group effects [103].

33.7.1.5 Taping One small study (n 5 20) [104] has reported a hallux taping technique in combination with exercises to be more effective than exercise alone. Both groups reported improvements in pain and walking ability after eight weeks, but greater improvements were noted in the group using the combined intervention.

33.7.1.6 Exercise Kim et al. [53] investigated the effects of the toe spread exercise in young adults (aged 19 29 years) with HV, in terms of HV angle and cross-sectional area of the abductor hallucis muscle. While the group using orthoses alone (n 5 10) experienced no improvement in these outcome measures, the group using orthoses in combination with the toe spread exercise (n 5 10) showed improved outcomes at eight weeks follow-up.

33.7.1.7 Botulinum toxin A injection One study has described the use of botulinum toxin A injections for HV [105]. This randomized controlled trial included 26 adults with painful HV. The treatment group (n 5 12) had injections of botulinum toxin type A into the adductor hallucis, flexor hallucis brevis, and extensor hallucis longus muscles, while the control group had normal saline injections into the same muscular sites. While both groups reported improved outcomes at one month, there was a significantly greater improvement in pain at two to three months in the treatment group compared to the control group. Box 33.3 below summarises the key points regarding effectiveness of nonsurgical treatments for hallux valgus.

33.7.2 Surgical treatment More than 150 different surgical procedures have been described in the management of HV [116]. The choice of surgical technique will depend upon radiographic assessment of the stage of severity of HV [72]. In consideration of surgery, the following risks must be considered and discussed with the patient: postoperative pain, risk of recurrence, as well as the general risks of surgery and anesthesia [48]. Fraissler et al. [117] report a higher rate of recurrence in patients with ligamentous laxity. Broad categories of surgical procedures include distal soft tissue procedures, first metatarsal osteotomies, osteotomies of the proximal phalanx of the hallux, arthrodesis, or resection arthroplasty [48]. There are many different types of first metatarsal osteotomies, and limited and inconsistent evidence surrounding which techniques achieve the best outcomes [101]. There is a consensus among experts that moderate HV deformities may be treated by distal metatarsal osteotomies, while severe HV deformities often require a proximal metatarsal osteotomy procedure. One systematic review by Ferrari et al. [106] has synthesized the evidence surrounding surgical HV treatments. This review included 11 randomized controlled trials investigating surgical options for HV, and in summary the authors reported a moderate level of evidence in favor of distal chevron osteotomy in improving pain and function, compared to no treatment. There is, however, only low-level evidence comparing different types of osteotomies, or comparing osteotomies with minimally invasive techniques. Two other systematic reviews have summarized the evidence surrounding percutaneous or minimally invasive procedures for HV [116,118]. Both reviews reported that the majority of evidence is low-level based largely on case series reports, and more high-quality studies are needed.

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33.8

541

Future directions for research

There are several areas requiring further research into this complex musculoskeletal condition. As causation cannot be inferred from cross-sectional studies, long-term prospective studies investigating the risk factors for the development of HV are needed. Second, mechanisms linking HV to postural instability and increased falls risk in older adults need to be further explored. Such studies should explore the links between changes in foot muscle morphology and neuromotor control, hallux flexor strength, postural sway, and altered gait patterns in HV. In terms of treatment pathways, good quality randomized controlled trials are needed investigating the effects of interventions in HV. There is currently a very limited evidence base for nonsurgical HV treatments, and the quality of evidence is low for surgical interventions. In addition to pain relief as a goal of treatment, nonsurgical interventions should focus on some of the identified functional deficits associated with HV, such as intrinsic muscle weakness and postural instability. Intervention studies should focus on pain and functional outcomes rather than changes in HV angle alone. Finally, randomized controlled trials must be well designed, with randomization, blinding of participants and examiners, as well as valid and reliable outcome measures. Outcomes from this future research will have a significant impact on the community given the prevalence of HV and the detrimental impact of falls and functional impairment in older adults.

33.9

Summary

HV is a common musculoskeletal condition with a complex and multifactorial etiology. Individuals with HV may experience associated foot pain, and moderate to severe deformity may be associated with muscle weakness, altered gait patterns, and impaired balance. Assessment of HV should include a thorough history, use of diagnostic imaging as relevant and functional and biomechanical assessments. Nonsurgical or surgical treatment options may then be tailored to the individual’s presenting concerns and functional deficits, whether that be foot pain, progression of deformity, difficulty fitting footwear, muscle weakness or problems with balance. More research is needed to understand the complex threedimensional biomechanics underlying HV, as well as to investigate the efficacy of the wide range of treatment options.

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Epidemiology of shoe wearing patterns over time in older women: associations with foot pain and hallux valgus. J Gerontol Ser A: Biol Sci Med Sci 2016;71(12):1682 7. [42] Garrow AP, Papageorgiou A, Silman AJ, Thomas E, Jayson MI, Macfarlane GJ. The grading of hallux valgus. The Manchester scale. J Am Podiatr Med Assoc 2001;91(2):74 8. [43] Nix S, Russell T, Vicenzino B, Smith M. Validity and reliability of hallux valgus angle measured on digital photographs. J Orthop Sports Phys Ther 2012;42(7):642 8. [44] Hopson MM, McPoil TG, Cornwall MW. Motion of the first metatarsophalangeal joint: reliability and validity of four measurement techniques. J Am Podiatr Med Assoc 1995;85:198 204. [45] Srivastava S, Chockalingam N, El Fakhri T. Radiographic measurements of hallux angles: a review of current techniques. Foot 2010;20 (1):27 31. [46] Coughlin MJ, Saltzman CL, Nunley JA. Angular measurements in the evaluation of hallux valgus deformities: a report of the ad hoc committee of the American Orthopaedic Foot & Ankle Society on angular measurements. Foot Ankle Int 2002;23(1):68 74. [47] Welck MJ, Al-Khudairi N. Imaging of hallux valgus: how to approach the deformity. Foot Ankle Clin 2018;23(2):183 92. [48] Hecht PJ, Lin TJ. Hallux valgus. Med Clin North Am 2014;98(2):227 32. [49] Menz HB, Munteanu SE, Landorf KB, Zammit GV, Cicuttini FM. Radiographic classification of osteoarthritis in commonly affected joints of the foot. Osteoarthr Cartil 2007;15(11):1333 8. [50] Matsubara K, Matsushita T, Tashiro Y, Tasaka S, Sonoda T, Nakayama Y, et al. Repeatability and agreement of ultrasonography with computed tomography for evaluating forefoot structure in the coronal plane. J Foot Ankle Res 2017;10 17-17. [51] Mickle KJ, Nester CJ. Morphology of the toe flexor muscles in older adults with toe deformities. Arthritis Care Res 2018;70(6):902 7.

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[52] Cameron AF, Rome K, Hing WA. Ultrasound evaluation of the abductor hallucis muscle: reliability study. J Foot Ankle Res 2008;1(1) 12-12. [53] Kim M-H, Yi C-H, Weon J-H, Cynn H-S, Jung D-Y, Kwon O-Y. Effect of toe-spread-out exercise on hallux valgus angle and cross-sectional area of abductor hallucis muscle in subjects with hallux valgus. J Phys Ther Sci 2015;27(4):1019 22. [54] Welck MJ, Singh D, Cullen N, Goldberg A. Evaluation of the 1st metatarso-sesamoid joint using standing CT—the Stanmore classification. Foot Ankle Surg 2018;24(4):314 19. [55] Collan L, Kankare JA, Mattila K. The biomechanics of the first metatarsal bone in hallux valgus: a preliminary study utilizing a weight bearing extremity CT. Foot Ankle Surg 2013;19(3):155 61. [56] Geng X, Wang C, Ma X, Wang X, Huang J, Zhang C, et al. Mobility of the first metatarsal-cuneiform joint in patients with and without hallux valgus: in vivo three-dimensional analysis using computerized tomography scan. J Orthop Surg Res 2015;10(1). [57] Kimura T, Kubota M, Taguchi T, Suzuki N, Hattori A, Marumo K. Evaluation of first-ray mobility in patients with hallux valgus using weightbearing CT and a 3-D analysis system: a comparison with normal feet. J Bone Jt Surg 2017;99(3):247 55. [58] Schweitzer ME, Maheshwari S, Shabshin N. Hallux valgus and hallux rigidus: MRI findings. Clin Imaging 1999;23(6):397 402. [59] Sanders AP, Weijers RE, Snijders CJ, Schon LC. Three-dimensional reconstruction of magnetic resonance images of a displaced flexor hallucis longus tendon in hallux valgus. J Am Podiatr Med Assoc 2005;95(4):401 4. [60] Kurihara T, Yamauchi J, Otsuka M, Tottori N, Hashimoto T, Isaka T. Maximum toe flexor muscle strength and quantitative analysis of human plantar intrinsic and extrinsic muscles by a magnetic resonance imaging technique. J Foot Ankle Res 2014;7 26-26. [61] Glasoe WM, Pena FA, Phadke V. Cardan angle rotation sequence effects on first-metatarsophalangeal joint kinematics: implications for measuring hallux valgus deformity. J Foot Ankle Res 2014;7 29-29. [62] Glasoe WM, Phadke V, Pena FA, Nuckley DJ, Ludewig PM. An image-based gait simulation study of tarsal kinematics in women with hallux valgus. Phys Ther 2013;93(11):1551 62. [63] Hurn SE, Vicenzino BT, Smith MD. Non-surgical treatment of hallux valgus: a current practice survey of Australian podiatrists. J Foot Ankle Res 2016;9(1):16. [64] Badlissi F, Dunn JE, Link CL, Keysor JJ, McKinlay JB, Felson DT. Foot musculoskeletal disorders, pain, and foot-related functional limitation in older persons. J Am Geriatr Soc 2005;53(6):1029 33. [65] Chaiwanichsiri D, Janchai S, Tantisiriwat N. Foot disorders and falls in older persons. Gerontology 2009;55(3):296 302. [66] Keysor JJ, Dunn JE, Link CL, Badlissi F, Felson DT. Are foot disorders associated with functional limitation and disability among communitydwelling older adults? J Aging Health 2005;17(6):734 52. [67] Leveille SG, Guralnik JM, Ferrucci L, Hirsch R, Simonsick E, Hochberg MC. Foot pain and disability in older women. Am J Epidemiol 1998;148(7):657 65. [68] Benvenuti F, Ferrucci L, Guralnik JM, Gangemi S, Baroni A. Foot pain and disability in older persons: an epidemiologic survey. J Am Geriatr Soc 1995;43(5):479 84. [69] Menz HB, Morris ME. Determinants of disabling foot pain in retirement village residents. J Am Podiatr Med Assoc 2005;95(6):573 9. [70] Menz HB, Roddy E, Thomas E, Croft PR. Impact of hallux valgus severity on general and foot-specific health-related quality of life. Arthritis Care Res (Hoboken) 2011;63(3):396 404. [71] Roddy E, Zhang W, Doherty M. Prevalence and associations of hallux valgus in a primary care population. Arthritis Care Res (Hoboken) 2008;59(6):857 62. [72] Vanore JV, Christensen JC, Kravitz SR, Schuberth JM, Thomas JL, Weil LS, et al. Diagnosis and treatment of first metatarsophalangeal joint disorders. section 1: hallux valgus. J Foot Ankle Surg 2003;42(3):112 23. [73] Abhishek A, Roddy E, Zhang W, Doherty M. Are hallux valgus and big toe pain associated with impaired quality of life? A cross-sectional study. Osteoarthr Cartil 2010;18(7):923 6. [74] Koski K, Luukinen H, Laippala P, Kivela SL. Physiological factors and medications as predictors of injurious falls by elderly people: a prospective population-based study. Age Ageing 1996;25(1):29 38. [75] Menz HB, Lord SR. The contribution of foot problems to mobility impairment and falls in community-dwelling older people. J Am Geriatr Soc 2001;49(12):1651 6. [76] Menz HB, Morris ME, Lord SR. Foot and ankle risk factors for falls in older people: a prospective study. J Gerontology Ser A: Biol Sci Med Sci 2006;61(8):866 70. [77] Mickle KJ, Munro BJ, Lord SR, Menz HB, Steele JR. ISB Clinical Biomechanics Award 2009. Toe weakness and deformity increase the risk of falls in older people. Clin Biomech (Bristol, Avon) 2009;24(10):787 91. [78] Tinetti ME, Speechley M, Ginter SF. Risk factors for falls among elderly persons living in the community. N Engl J Med 1988;319 (26):1701 7. [79] Menz HB, Lord SR. Foot pain impairs balance and functional ability in community-dwelling older people. J Am Podiatr Med Assoc 2001;91 (5):222 9. [80] Spink MJ, Fotoohabadi MR, Wee E, Hill KD, Lord SR, Menz HB. Foot and ankle strength, range of motion, posture, and deformity are associated with balance and functional ability in older adults. Arch Phys Med Rehab 2011;92(1):68 75. [81] Hurn SE, Vicenzino B, Smith MD. Functional impairments characterizing mild, moderate, and severe hallux valgus. Arthritis Care Res 2015;67 (1):80 8. [82] Menz HB, Morris ME, Lord SR. Foot and ankle characteristics associated with impaired balance and functional ability in older people. 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[83] Mickle KJ, Munro BJ, Lord SR, Menz HB, Steele JR. Gait, balance and plantar pressures in older people with toe deformities. Gait Posture 2011;34(3):347 51. [84] Zhang X, Schu¨tte KH, Vanwanseele B. Foot muscle morphology is related to center of pressure sway and control mechanisms during singleleg standing. Gait Posture 2017;57:52 6. [85] Endo M, Ashton-Miller JA, Alexander NB. Effects of age and gender on toe flexor muscle strength. J Gerontol Ser A: Biol Sci Med Sci 2002;57(6):M392 7. [86] Sanders AP, Snijders CJ, van Linge B. Medial deviation of the first metatarsal head as a result of flexion forces in hallux valgus. Foot Ankle 1992;13(9):515 22. [87] Mortka K, Lisi´nski P, Wiertel-Krawczuk A. The study of surface electromyography used for the assessment of abductor hallucis muscle activity in patients with hallux valgus. Physiother Theory Pract 2018;34(11):846 51. [88] Kelly LA, Kuitunen S, Racinais S, Cresswell AG. Recruitment of the plantar intrinsic foot muscles with increasing postural demand. Clin Biomech (Bristol, Avon) 2012;27(1):46 51. [89] Nix S, Vicenzino B, Collins N, Smith M. Gait parameters associated with hallux valgus: a systematic review. J Foot Ankle Res 2013;6(1):9. [90] Deschamps K, Birch I, Desloovere K, Matricali GA. The impact of hallux valgus on foot kinematics: a cross-sectional, comparative study. Gait Posture 2010;32(1):102 6. [91] Canseco K, Rankine L, Long J, Smedberg T, Marks RM, Harris GF. Motion of the multisegmental foot in hallux valgus. Foot Ankle Int 2010;31(2):146 52. [92] Hwang SJ, Choi HS, Cha SD, Lee KT, Kim YH, Multi-segment foot motion analysis on hallux valgus patients, In: Proceedings of the 2005 IEEE Engineering in Medicine and Biology 27th Annual Conference, September 1 4, 2005, Shanghai, China. [93] Glasoe WM, Jensen DD, Kampa BB, Karg LK, Krych AR, Pena FA, et al. First ray kinematics in women with rheumatoid arthritis and bunion deformity: a gait simulation imaging study. Arthritis Care Res 2014;66(6):837 43. [94] Bryant A, Tinley P, Singer K. Radiographic measurements and plantar pressure distribution in normal, hallux valgus and hallux limitus feet. Foot 2000;10(1):18 22. [95] Martı´nez-Nova A, Sa´nchez-Rodrı´guez R, Pe´rez-Soriano P, Llana-Belloch S, Leal-Muro A, Pedrera-Zamorano JD. Plantar pressures determinants in mild hallux valgus. Gait Posture 2010;32(3):425 7. [96] Kadono K, Tanaka Y, Sakamoto T, Akiyama K, Komeda T, Taniguchi A, et al. Plantar pressure distribution under the forefeet with hallux valgus during walking. J Nara Med Assoc 2003;54(5 6):273 81. [97] Menz HB, Morris ME. Clinical determinants of plantar forces and pressures during walking in older people. Gait Posture 2006;24(2):229 36. [98] Mueller MJ, Hastings M, Commean PK, Smith KE, Pilgram TK, Robertson D, et al. 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Manual and manipulative therapy compared to night splint for symptomatic hallux abducto valgus: an exploratory randomised clinical trial. Foot 2011;21(2):71 8. [104] Bayar B, Erel S, Simsek I. The effects of taping and foot exercises on patients with hallux valgus: a preliminary study. Turk J Med Sci 2011;41:403 9. [105] Wu KP-H, Chen C-K, Lin S-C, Pei Y-C, Lin R-H, Tsai W-C, et al. Botulinum Toxin type A injections for patients with painful hallux valgus: a double-blind, randomized controlled study. Clin Neurol Neurosurg 2015;129:S58 62. [106] Ferrari J. Hallux valgus (bunions). BMJ Clin Evid 2014;2014:1112. [107] Kilmartin TE, Barrington RL, Wallace WA. A controlled prospective trial of a foot orthosis for juvenile hallux valgus. J Bone Jt Surg Br Vol 1994;76(2):210 14. [108] Reina M, Lafuente G, Munuera PV. Effect of custom-made foot orthoses in female hallux valgus after one-year follow up. Prosthet Orthot Int 2013;37(2):113 19. 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Chapter 34

Osteoarthritis of the Foot and Ankle Kade L. Paterson1, Luke A. Kelly2 and Michelle D. Smith3 1

Centre for Health, Exercise and Sports Medicine, Department of Physiotherapy, School of Health Sciences, Faculty of Medicine Dentistry & Health

Sciences, The University of Melbourne, Melbourne, VIC, Australia, 2School of Human Movement and Nutrition Sciences, The University of Queensland, St Lucia, QLD, Australia, 3School of Health & Rehabilitation Sciences, The University of Queensland, St Lucia, QLD, Australia

Abstract This chapter provides an overview of osteoarthritis (OA) from an epidemiological, clinical, pathological, and biomechanical perspective. While the chapter has a clinical focus, the primary objective is to present information on the biomechanical aspects of OA affecting the first metatarsophalangeal joint, the midfoot, and the ankle. This is complemented by current knowledge of the different OA risk factors and the biomechanical and clinical effects of interventions for foot and ankle OA. The chapter concludes by highlighting the areas for future research in the field.

34.1

Introduction

Osteoarthritis (OA) is the most common joint disorder, and the primary symptom of OA is pain. The disease is estimated to affect around 10% of men and 18% of women over the age of 60 years worldwide [1], and its prevalence is expected to dramatically increase with an aging population and growing obesity rates. Prevalence of OA in the weightbearing joints has historically been considered to be highest at the knee; however, a recent large epidemiology study has reported the population prevalence of symptomatic radiographic foot OA to be 16.7%, making it as common as knee OA [2]. Pain due to OA leads to significant physical disability, which in turn causes problems with mobility and other daily activities, and significantly reduces quality of life [3,4]. The condition also causes a substantial economic burden, with recent estimates of the indirect and direct costs due to musculoskeletal conditions such as OA ranging from 0.7% to 2.5% of gross national product in western countries [1].

34.1.1 Osteoarthritis symptoms and diagnosis In addition to pain, cardinal symptoms of OA include transient morning stiffness, crepitus (a grating or cracking sound or sensation experienced when the joint is moved), bony tenderness, and bony enlargements. Accordingly, leading international OA organizations now suggest that the condition can be diagnosed based on a combination of these clinical symptoms, in addition to age. For instance, the National Institute for Health and Care Excellence in the United Kingdom states that OA can be diagnosed clinically without further investigations if the patient is aged 45 years or over, has activity related joint pain, and either has no morning stiffness, or has morning stiffness that lasts less than 30 minutes [5]. Unfortunately, much of the research that informed these guidelines was based on patients with knee and hip OA; therefore, it is unknown whether OA of the foot or ankle can be diagnosed using these criteria. Regarding imaging, the latest recommendations from an expert international OA body (the European League Against Rheumatism) suggest that radiography or other imaging is not required to make a diagnosis of OA in patients with a typical presentation of the disease or to monitor disease progression [6]. The guidelines did acknowledge that imaging may be used in atypical cases, or if there is a rapid and unexpected progression of symptoms. However, it is not clear whether this advice is appropriate for foot and/or ankle OA given the recommendations are mostly based on hand, knee, and hip OA. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00026-3 © 2023 Elsevier Inc. All rights reserved.

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34.1.2 Structural changes The main OA-related joint changes identified on X-ray include osteophyte formation, joint space narrowing, subchondral sclerosis, and cysts (Fig. 34.1). Furthermore, there are several grading systems that use these features to classify OA [8 10]. However, with the advent of advanced imaging modalities (such as magnetic resonance imaging), it is now acknowledged that OA is a complex disorder affecting the whole joint. Specifically, common structural signs identified in OA include changes to joint cartilage, subchondral bone, joint ligaments, the joint capsule, the synovial membrane, and local and systemic inflammation.

Bone Bone cysts Joint capsule

Cartilage

Cartilage fragments

Sclerotic bone

Early osteoarthritis

Normal

Irregular joint space Fragmented cartilage Loss of cartilage Sclerotic bone Cystic change

Periarticular fibrosis

Osteophytes

Subchondral cysts

Osteophytes

Calcified cartilage

Moderate osteoarthritis Osteophytes Periarticular fibrosis Calcified cartilage

Advanced osteoarthritis Loss of joint space Large osteophytes Large subchondral cysts Morphological change

FIGURE 34.1 Structural joint changes seen in the development and progression of knee osteoarthritis. From Cividino, A. and J. O’Neill, Osteoarthritis, in Essential imaging in rheumatology, J. O’Neill, Editor. 2015, Springer New York: New York. p. 259 277 [7].

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34.1.3 Risk factors and classification of osteoarthritis In most instances, OA is recognized as being due to a combination of systemic risk factors occurring in the context of an altered mechanical environment. Leading systemic OA risk factors include increased age, obesity, female gender, reduced activity levels, and genetic susceptibility [11]. Mechanical risk factors that have been identified include abnormal joint morphology and alignment, altered joint loading, muscle weakness, and joint injury [11]. Historically, these have also been used to classify OA as primary or secondary. Primary or idiopathic OA is attributed to obesity, aging, and genetic susceptibility [12]. Secondary OA is due to external factors such as joint injury, trauma, surgery, or congenital joint abnormalities, amongst others [12]. While the relative contribution of these and other risk factors varies by individual and OA subtype, throughout this chapter we will largely focus on the biomechanical factors associated with OA and its treatment.

34.1.4 Foot and ankle osteoarthritis subtypes Foot OA is often considered a multi-joint disease because the median number of joints affected in people aged over 65 is four [13]. Nonetheless, distinct subtypes of foot and ankle OA have now been identified. Specifically, a recent large population-based study has shown that two distinct foot OA phenotypes exist; a polyarticular form that largely affects the midfoot joints, and isolated OA of the first metatarsophalangeal joint (MTPJ) [14]. Furthermore, ankle OA can also be considered as a separate entity, largely due to the different anatomy and biomechanical properties of the joint [15]. The following sections provide an overview of OA of the first MTPJ, midfoot, and ankle. For each region, we discuss the etiology, diagnosis, and effects of treatment, with a particular emphasis on relevant biomechanical factors.

34.2

First metatarsophalangeal joint osteoarthritis

34.2.1 Etiology and impact The first metatarsophalangeal (MTP) joint is the foot joint most commonly affected by OA. The prevalence rate of first symptomatic radiographic MTPJ OA is 7.8% of people aged over 50 years [2], which is slightly higher than prevalence rates for hip OA [16,17]. Consequently, much of the research investigating the etiology, impact, and treatment of foot OA focuses on the first MTPJ. Close to three quarters (69%) of patients describe the pain due to first MTPJ OA as disabling [2]. Bergin et al. [18] reported that people with first MTPJ OA have greater difficulties performing functional weight-bearing activities, finding suitable footwear, and that they perceive their feet to be in a poorer state of health than those without the condition. People with first MTPJ pain score lower on the overall health satisfaction and psychological domains of the World Health Organization Quality of Life questionnaire [19]. First MTPJ OA symptoms have remained relatively stable over time in the absence of treatment, with pain levels, reduced function and quality of life, and depressive symptoms showing very little change over an 18-month period [20]. Changes in symptoms over longer time periods are currently unknown; however, progression of first MTPJ radiographic OA and osteophyte development, in particular, have been found over a 19-year period in women from the Chingford 1000-women study in the United Kingdom [21].

34.2.2 Clinical findings Clinically, first MTPJ OA often presents as localized pain and joint stiffness, palpable dorsal bony enlargements with associated tenderness, and reduced joint range of motion. Despite these established clinical signs; however, there are currently no accepted clinical criteria for diagnosing first MTPJ OA. In the absence of clinical diagnostic criteria for first MTPJ OA, a large international survey suggested there were six common clinical assessments for patients with first MTPJ OA [22,23]. These included the following: (1) pain on walking over the past week, (2) first MTPJ range of motion, (3) ankle joint range of motion, (4) foot posture (foot posture index), (5) resting calcaneal stance position, and (6) palpation to determine pain location. However, it should be highlighted that the aim of this exercise is not to advise or recommend clinical diagnostic criteria, but rather to document how the condition is currently assessed in clinical practice. Other researchers have attempted to associate clinical findings with the presence of radiographic OA at the first MTPJ. For instance, Zammit and colleagues developed a diagnostic rule for first MTPJ OA in which clinical observations were used to predict the presence of radiographic first MTPJ OA in patients with first MTPJ pain [24]. The five

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signs associated with radiographic OA of the first MTPJ were pain duration greater than 25 months, the presence of a dorsal exostosis, hard-end feel, crepitus, and less than 64 of first MTPJ dorsiflexion. Using these signs, a cut off score of $ 3 resulted in a sensitivity of 88%, specificity of 71%, and accuracy of 84%. Another study found that radiographic OA severity at the first MTPJ was associated with an increased prevalence of dorsal hallux and first MTPJ pain, hallux valgus, first interphalangeal joint hyperextension, keratotic lesions on the dorsal aspect of the hallux and first MTPJ, decreased first MTPJ dorsiflexion, ankle/subtalar joint eversion and ankle joint dorsiflexion range of motion, and a more everted foot posture [25].

34.2.3 Structural and biomechanical features Several studies have identified structural changes present in people with first MTPJ OA compared to those without the disease. In a systematic review, Zammit and colleagues [26] found that first MTPJ OA patients had a dorsiflexed first compared to second metatarsal, a plantarflexed forefoot relative to the hindfoot, reduced first MTPJ range of motion, a longer proximal and distal phalanx and a longer medial and lateral sesamoid, and a wider first metatarsal and proximal phalanx. Measures of static foot posture and arch height were not different between those with and without first MTPJ OA. Many of the altered structural features identified in people with first MTPJ OA may have predisposed these individuals to the disease, rather than been a consequence of OA. According to this theory, the dorsiflexed position of the first metatarsal, longer lengths of the proximal and distal phalanx and sesamoids, and the wider first metatarsal and proximal phalanx, restrict first metatarsal plantarflexion during walking, which in turn limits the ability of the first metatarsal to glide and rotate on the fixed proximal phalanx during propulsion. This causes the proximal phalanx to “jam” on to the dorsal aspect of the first metatarsal head, increasing dorsal joint compressive forces which can cause damage to cartilage and subchondral bone, and ultimately lead to dorsal osteophytes, joint space narrowing, and pain, in people with first MTPJ OA [26 28]. Although none of the studies included in the systematic review were prospective, (hence it is not possible to determine whether these altered structural factors were causative or a consequence of the disease), the fact that studies such as the one by Munuera et al. [28] included relatively young participants with early stage disease may suggest that the structural radiographic features preceded symptomatic and radiographic signs of OA. In addition to the structural changes observed in people with first MTPJ OA, several studies have provided evidence of biomechanical alterations during walking. Most notably, patients with first MTPJ OA walk with reduced sagittal plane range of motion at the joint and reduced overall dorsiflexion (Fig. 34.2) [29]. This interferes with effective weight transfer during the propulsive phase of gait and results in an apropulsive gait pattern characterized by compensatory increased knee and hip flexion [30]. In fact, compensations such as these have been suggested to potentially increase stress on more proximal joints such as the knee, which may partly explain why foot pain is a risk factor for developing knee OA in the future [31], and for future worsening of knee pain in people with existing knee OA [32].

Hallux: Sagittal (+) Dorsi / (-) Plantar

Angle [deg]

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% Gait Cycle FIGURE 34.2 Average hallux dorsiflexion and plantarflexion during complete gait cycle (Hallux Rigidus group vs Control group). Circles denote phases of gait cycle with significantly different positions. Note: Hallux kinematics are calculated relative to forefoot. From Canseco, K., et al., Quantitative characterization of gait kinematics in patients with hallux rigidus using the Milwaukee Foot Model. J Orthopaedic Res, 2008. 26(4): p. 419 427 [29].

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FIGURE 34.3 Representation of mechanical overloading and antalgic patterns of load distribution. On the left, the gray ovals indicate the regions assessed: hallux, second, fourth, and fifth metatarsals. On the right, mechanical overloading (evidenced as increased loading on hallux) is depicted in the top image, and an antalgic strategy (evidenced as increased loading on second, fourth, or fifth metatarsal) is depicted in the bottom image. From Rao, S., et al., Are pressure time integral and cumulative plantar stress related to first metatarsophalangeal joint pain? Results from a communitybased study. Arthritis Care Res, 2016. 68(9): p. 1232 1238 [33].

Plantar pressures have been shown to be altered in people with first MTPJ OA, and there are two conflicting theories regarding these [33]. Specifically, the reduced joint range of motion observed in first MTPJ OA may increase plantar pressures under the hallux, which may contribute to increased pain at the joint. Alternatively, people with the disease may move their center of pressure more laterally under the forefoot during walking to avoid the painful joint (Fig. 34.3). In fact, there is some evidence for both strategies. Several studies have found greater force [34] and peak pressures [34 36] under the hallux in people with first MTPJ OA compared to controls, albeit the Zammit study [34] was comprised of asymptomatic people with radiographic first MTPJ OA. In contrast, other researchers have found that people with a painful first MTPJ walk with lower hallux pressures [33], and a more lateral center of pressure, both of which imply a pain avoidance strategy. While it is not immediately clear why these studies have conflicting findings, the differing inclusion and exclusion criteria (e.g., ages, symptom duration, etc.), varying case definitions of first MTPJ OA/hallux rigidus (e.g., radiographic, symptomatic or a combination), different data collection protocols (e.g., number of trials and steps, plantar pressure mat or platform), and outcome measures (e.g., peak plantar pressure, center of pressure, pressure time integral) may have contributed.

34.2.4 Clinical and biomechanical effects of conservative treatment Although first MTPJ OA is as common as hip OA, high-quality clinical research on conservative treatment options is lacking. In fact, there have only been two non-pharmacological, non-surgical randomized clinical trials (RCTs) published on OA of the first MTPJ [37,38], and neither included a control group. In addition to these, there is also one trial comparing hyaluronic acid vs saline injections [39], one comparing hyaluronic acid vs steroid injections [40], a RCT feasibility study [41], and a protocol study for an upcoming RCT comparing shoe stiffening inserts to foot orthoses [42]. Of the two non-pharmacological, non-surgical clinical trials, only one compared two different biomechanical interventions, namely rocker-soled shoes and foot orthoses with a cut out under the first MTPJ, in 102 participants with OA of the first MTPJ (Fig. 34.4) [38]. Clinical outcomes from this study showed both interventions improved the 12-week change in pain (measured using the subscale of the Foot Health Status Questionnaire); however, there was no statistical difference between the two groups. Despite this, there were several secondary findings that are clinically relevant. Most notably, the foot orthoses group had greater adherence, were less likely to experience adverse events, and were more likely to report global improvements in first MTPJ OA symptoms.

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FIGURE 34.4 Rocker-soled shoes and foot orthoses with a cut out under the first MTPJ used in the Menz et al. trial. Top: plantar surface of left foot orthosis. Bottom: dorsal surface of right foot orthosis. From Menz, H.B., et al., Rocker-sole footwear vs prefabricated foot orthoses for the treatment of pain associated with first metatarsophalangeal joint osteoarthritis: study protocol for a randomised trial. BMC Musculoskelet Disord, 2014. 15(1): p. 86 [43].

There is a biomechanical rationale for explaining how rocker-sole shoes and foot orthoses may improve pain associated with first MTPJ OA. Firstly, the curved sole of the rocker-sole shoe is designed to “rock” the foot from heel contact to toe off during gait, reducing the need for the first MTPJ to dorsiflex thus potentially decreasing joint compression and hallux plantar pressures [44]. To date, only one study has investigated the effects of rocker-sole shoes on first MTPJ kinematics, showing these shoes do reduce first MTPJ dorsiflexion as expected [45]. Evidence regarding the effects of these shoes on plantar pressures, however, is mixed. In a preliminary biomechanical study, Menz et al. [46] reported that plantar pressures were reduced under the first and lesser MTPJs, which supported earlier research showing two different rocker-sole designs reduced peak pressures under first MTPJ compared to control footwear [47]. However, although one of these two latter shoes also reduced peak pressure under the hallux, the other increased hallux plantar pressures [47], while Menz and colleagues found no change in hallux plantar pressures [46]. Foot orthoses are probably effective for first MTPJ OA by altering joint kinematics and plantar pressures. Specifically, foot orthoses with a cut out under the first MTPJ have been suggested to facilitate first metatarsal plantarflexion [48], allowing it to glide and rotate on the proximal phalanx and sesamoids during propulsion [49]. This is believed to reduce dorsal joint compressive forces [26 28] which in turn potentially reduces pain [49]. There is some limited evidence to support this theory. In a study of 46 asymptomatic women with restricted weightbearing first MTPJ motion, researchers found a greater plantarflexion angle of the first metatarsal when walking in a foot orthosis with a cut out under the first MTPJ [50]. Another small case series found that although a similarly designed foot orthoses improved pain over 24 weeks, the device did not increase joint dorsiflexion, and there was no relationship between joint range of motion and improvements in pain [51]. Like rocker-soled shoes, it is also possible that foot orthoses with a cut out under the first MTPJ may improve pain via a redistribution of plantar pressures. In the only study to investigate the effects of this type of foot orthoses on plantar pressures, Menz et al. found reduced pressures under the first MTPJ, and increased pressures under the lesser toes, while walking with the foot orthoses compared to footwear alone, in people with symptomatic first MTPJ OA [46]. This suggests that these devices shift pressures more laterally away from the painful region during propulsion.

34.3

Midfoot osteoarthritis

34.3.1 Etiology and impact Polyarticular foot OA is a distinct phenotype of OA that commonly affects the joints of the midfoot [52,53]. Midfoot OA is rarely observed as an isolated entity, often affecting the first MTPJ concurrently [52]. For the purposes of this section, polyarticular OA will be referred to as midfoot OA. Symptomatic radiographic OA of the midfoot most commonly affects the cuneiometatarsal (CMT) joints (first CMT 3.9%, second CMT 6.8%), the naviculocuneiform (NC, 5.2%) joint and the talonavicular (TN, 5.8%) joint. The combined prevalence of midfoot OA, when considering any one of these joints, is 12% [2,54]. The prevalence of this condition generally increases with age, especially in females. A higher body mass index (BMI) and a lower socioeconomic status is also associated with increased prevalence [2,54]. Pain is described as disabling in three-quarters of people with midfoot OA [2,54], who report difficulty with tasks of daily living and reductions in health related quality of life [52,53]. Similar to first MTPJ OA, pain, quality of life, and functional impairments associated with midfoot OA appear to be persistent, but stable over a period of 18 months in the absence of treatment [20]. However, little is known about the long-term progression of this condition.

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34.3.2 Clinical findings People with midfoot OA often report the hallmark signs of OA in other joints, such as localized pain and tenderness, joint stiffness, and bony enlargements [55,56]. Owing to the polyarticular nature of midfoot OA, the precise location of pain can vary, depending on the number and location of midfoot joints involved. However, pain can generally be localized to the dorsal and/or medial aspects of the midfoot. Pain is often described as worsening with tasks requiring greater propulsive forces to be generated, such as stair ascent [57]. Similar to OA of the first MTPJ, there is currently no validated clinical diagnostic tool for midfoot OA. Thomas et al. [56] studied a large population of people with midfoot pain to develop a diagnostic model based on a brief clinical assessments and person level data such as age, gender, and BMI. Predictors were included in the model based on three criteria: (i) known risk factors for symptomatic OA at other joints, (ii) a biomechanical link to symptomatic midfoot OA, and (iii) be applicable in a clinical setting. A total of 16 predictor variables were included across three categories, including static foot posture, joint range of motion, and palpation and observation [56]. Despite 10 of the predictor variables being associated with the presence of midfoot OA, these associations were mostly too weak to be included in the model, particularly after adjusting for other covariates. The strongest individual predictor of midfoot OA was arch index, with a more everted foot evident in those with symptomatic midfoot OA. The primary finding of this study, however, is that person level information, such as age, gender, and BMI provide only provide marginal diagnostic information for midfoot OA [56]. Furthermore, the addition of a brief clinical assessment adds little further diagnostic value. Despite these conclusions, however, the poor predictive capacity may be explained by several factors, and most notably, the lack of specificity of many of the clinical assessments to the joints of the midfoot, or relevance to midfoot OA symptoms. For example, the joint range of motion variables do not directly evaluate the midfoot, but rather joints distal and proximal to this region of the foot. Likewise, pain due to midfoot OA is largely dorsal and/or medial not near the plantar fascia insertion or midpoint of the medial longitudinal arch. Additionally, the model did not include measures to evaluate soft tissue pathology that may help to further discriminate between midfoot OA and other sources of midfoot pain. Thus, it is currently unclear whether clinical signs may be used to successfully diagnose midfoot OA.

34.3.3 Structural and biomechanical features Despite the relatively common occurrence and the disabling nature of midfoot OA, remarkably little is known about the underlying structural and biomechanical features associated with the onset and progression of this condition. This lack of understanding is due to a combination of factors, but perhaps the greatest barrier is our ability to accurately quantify the biomechanical function of the multiple articulations in the midfoot in vivo. Furthermore, midfoot OA is polyarticular in nature, with between-person variability in the number and location of joints involved [52,53]. This may indicate multiple combinations of structural and biomechanical factors contributing to the onset of this condition. A recently published systematic review by Lithgow and colleagues [58] reported differences in foot structure and lower limb function between individuals with and without midfoot OA. Based on the data from ten included studies, they found individuals with midfoot OA had a more everted foot posture, longer central metatarsals, greater first ray mobility, reduced subtalar and first MTPJ range of motion, and altered midfoot plantar pressures. However, all the included data in this review were obtained from cross-sectional studies, so little can be inferred regarding the role of these features in the pathogenesis of midfoot OA. A consistent radiographic finding in people with midfoot OA is that the second metatarsal is longer than the first metatarsal [59,60]. While this observation is not uncommon in a general healthy population, the presence of a functionally longer second metatarsal could increase the leverage of the metatarsal about the second tarsometatarsal joint [61], potentially increasing the external moment and dorsal joint contact forces at the tarsometatarsal joints. While there is no prospective data to confirm the theory, it seems plausible that this relationship might be causal in nature, given that metatarsal length is unlikely to change during the course of disease progression. Another consistent structural feature observed with both radiographic and clinical assessments, is the presence of a more flatfoot posture [53,56,60,62,63]. It has been suggested that the combination of a flatfoot posture, lower medial longitudinal arch, and greater midfoot loading [62,63] in people with midfoot OA may contribute to increased dorsal compressive forces through the midfoot [62]. While this is yet to be shown in vivo, there is cadaveric data to suggest that greater foot eversion and increased axial loading increase the dorsal and central compressive forces through the talonavicular joint [64]. Despite the consistent observation of a more everted foot posture in people with midfoot OA, it needs to be highlighted that this finding could also potentially occur as a mechanical consequence of the condition. For example, a flatfoot posture may be a result of alterations in articular congruence, reduced joint ranges of motion, and muscle weakness, commonly observed in OA.

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People with midfoot OA display reductions in non-weight bearing, subtalar joint inversion and first MTPJ dorsiflexion range of motion [53]. The finding of reduced first MTPJ range of motion is not surprising, given the polyarticular nature of midfoot OA that commonly involves the first MTPJ in addition to joints of the midfoot [52,53]. The reduced subtalar joint inversion suggests that this condition may also influence the function of proximal joints within the foot, and is also consistent with the reports of a more everted foot posture in people with midfoot OA. In addition to the common structural features observed in midfoot OA, several functional differences have also been observed. Rao and colleagues compared subtalar joint and first ray range of motion during walking and stair ascent in individuals with midfoot OA, compared to matched healthy controls [57]. They reported that individuals with midfoot OA had higher calcaneal eversion during stair ascent. This finding extends from the previously described finding of a more everted foot posture in midfoot OA [53] and further highlights the potentially important role of hindfoot eversion in midfoot OA across the entire foot. The relationship between first ray motion during walking and stair ascent is less clear. In the same study, Rao et al. [57] reported that first ray range of motion was reduced during walking (B6 degrees reduction), but similar during stair ascent in people with midfoot OA. The authors speculated that reduced first ray range of motion during walking, may be a strategy to guard or stiffen the midfoot to reduce joint compressive stress. During stair ascent, the torques generated across the ankle and midfoot are much greater and thus, this guarding strategy may not be possible. Their suggestion does fit with patient reported pain data from the same study, indicating that the stair ascent task elicited substantial pain. In a separate study, Rao et al. [65] also explored differences in regional foot plantar loading in people with midfoot OA, compared to healthy matched controls. They reported increased plantar pressures in the heel, midfoot and medial forefoot, in conjunction with higher ground contact times for the heel and midfoot. Menz et al. [62] also reported greater peak midfoot plantar pressures during walking in people with midfoot OA. It is possible that a more flatfoot posture and reduced subtalar joint motion in midfoot OA could result in a medial shift in plantar foot loading. The increase in ground contact time and greater proportion of time spent on the heel and midfoot, could be a strategy to reduce peak joint compression stress through the midfoot. A longer ground contact would allow the same impulse to be produced by the muscles of the lower limb, with reduced peak ground reaction forces. Further, prolonged time on the heel and midfoot would delay and reduce the forward shift of the center of pressure, potentially reducing the length of the external (ground reaction force) moment arm about the ankle and midfoot joints (Fig. 34.5).

FIGURE 34.5 Representative plantar pressure recordings of participants without (A) and with (B) medial midfoot OA. From Menz, H.B., et al., Foot structure and function in older people with radiographic osteoarthritis of the medial midfoot. Osteoarthr Cartil, 2010. 18(3): p. 317 322 [62].

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There has been very little research into the changes in lower limb neuromuscular function in people with midfoot OA. Yi et al. [66] reported electromyography data from the tibialis anterior, gastrocnemius medialis and soleus muscles during self-paced walking. They found no differences in the activation patterns of these muscles in people with midfoot OA compared to healthy control subjects. Despite the lack of difference observed in muscle activation patterns during walking, it is likely that some alterations in muscle force producing capacity may exist in people with symptomatic midfoot OA. A recent study by Arnold et al. [67] has found that people with symptomatic midfoot OA demonstrate substantial weakness in the foot and leg muscles compared to asymptomatic controls. These deficits were evident in all muscle groups evaluated, with maximal voluntary isometric force deficits in ankle dorsiflexion (219%), ankle plantarflexion (226%), foot inversion (230%), foot eversion (226%), digital flexion (22%), and hallux flexion (27%). Further research is required to explore how altered force production and neuromuscular control contribute to the onset and progression of midfoot OA.

34.3.4 Clinical and biomechanical effects of conservative treatment Despite the common occurrence and debilitating nature of midfoot OA, there is limited information about the efficacy of conservative treatment interventions. In a primary care setting, pharmacological interventions are the most commonly adopted management strategy [68]. Common non-pharmacological treatments include shoe stiffening inserts and arch supporting foot orthoses, albeit there is limited evidence to support either management approach. Two case series have been published, evaluating the short term (4 weeks) effect of shoe stiffening inserts on foot pain and function [66,69]. While a randomized feasibility trial provides some insight into the potential for arch supporting foot orthoses to deliver short—medium term (12 weeks) benefits in patient-reported pain and function. Unlike knee and hip OA, there are currently no studies evaluating the effectiveness of exercise interventions in the management of midfoot OA. Rao et al. [69] reported a case series of 30 participants with midfoot OA, who were provided shoe stiffening inserts made from carbon fiber, for use over a four week period. Another case series by Rao et al. [70] reported the effect of the same shoe stiffening inserts on pain and function in 20 females with midfoot OA. In both case series, kinematic and kinetic data were collected at the start of each study, to link pain and function outcomes to biomechanical alterations induced by the use of the shoe stiffening inserts. Using the Foot Function Index Revised (FFI-R) [71], significant improvements in pain and function were observed in both case series, with the magnitude of benefit ranging between 12% 20% in overall FFI-R scores. The reductions in pain were observed alongside reduced first MTPJ and first ray range of motion, and reductions in medial plantar loading. No change in ankle or subtalar joint kinematics were observed when participants walked with the shoe stiffening inserts, compared to their normal footwear. The authors suggested that the shoe stiffening inserts may provide a simple clinical tool to shift loading away from the medial aspect of the foot and reduce excessive first ray rotation, thus limiting articular stresses at the tarso-metatarsal joints. Further supporting this idea, the first case series reported a modest positive relationship between first ray range of motion and pain subscale scores from the FFI-R, with higher pain scores being evident in those with greater first ray range of motion during walking [69]. Furthermore, the second case series reported that reductions in pain and improvements in function were associated with greater reductions in medial plantar pressures [70]. Halstead et al. [72] conducted a double blind randomized feasibility study examining the clinical and biomechanical effects of arch supporting foot orthoses in the treatment of midfoot OA, as well as the feasibility of a larger RCT. With a relatively small cohort of 37 participants, 19 were allocated to the foot orthoses group and 18 to a sham insole group (Fig. 34.6). Participants in both groups reported improvements in pain and function, with the foot orthoses group showing a trend toward greater improvements. However, the magnitude of difference between the two groups was smaller than a pre-defined clinically meaningful margin, according to the Manchester Foot Pain and Disability Index [73]. The study also reported that when using foot orthoses, participants walked with the hindfoot in slightly more inversion, with an increase in midfoot plantar pressure, compared to the sham group. These findings are consistent with the biomechanical rationale for arch supporting foot orthoses use, with the aim to reduce foot eversion and shift load within the foot. However, the increase in medial force produced by foot orthoses use may seem somewhat counterintuitive, given the previously established link between increased medial plantar pressures and increased pain in people with midfoot OA [65,70]. The authors suggest that the increased force under the midfoot is an indication of the mechanical effect of the orthoses in providing an upwards reaction force at the midfoot, supporting the medial longitudinal arch and potentially reducing the compressive stress through these joints. There appears to be good potential and sound rationale for the use of foot orthoses in the management of symptoms associated with midfoot OA. However, fully powered RCTs are required before this can be undertaken with any confidence.

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FIGURE 34.6 Posterior-medial view of (A) functional foot orthoses and (B) sham orthoses used in the Halstead et al. trial. From Halstead, J., et al., Foot orthoses in the treatment of symptomatic midfoot osteoarthritis using clinical and biomechanical outcomes: a randomised feasibility study. Clin Rheumatol, 2016. 35(4): p. 987 996 [72].

34.4

Ankle osteoarthritis

34.4.1 Etiology and impact Consistent with OA at the first MTPJ and midfoot, there is limited information in the literature on the prevalence of ankle OA [15]. A 2018 clinical and radiographic evaluation of 557 community-dwelling adults identified that the prevalence of ankle pain was 11.7% and the prevalence of symptomatic radiographic ankle OA was 3.4% [15] However, this study only assessed for radiographic evidence of OA at the talus, tibia, and fibula, and did not consider the articulation between the talus and calcaneus. Thus, the prevalence of ankle OA (encompassing both the talocrural and subtalar joints) is expected to be even higher. Research suggests that the prevalence of ankle OA may be higher in certain population groups. A study found that 97% of football players and 88% of ballet dancers had radiographic ankle OA, and this increased to 100% for athletes from both sports who were over 25 years of age and had been participating for over 8 years [74]. The higher prevalence of ankle OA in sporting populations is likely due to the fact that the disease is commonly associated with one or multiple previous injuries to the ankle, such as ankle sprains or fractures. In fact, it is estimated that approximately 80% of ankle OA cases are post-traumatic in nature [75,76]. As ankle sprains are one of the most common injuries seen in emergency departments and in sporting populations [77], it is not surprising that post-traumatic OA occurs more commonly at the ankle than other joints of the lower limb [76]. The most common age to sustain an ankle sprain is between 10 and 19 years [77], and the mean latency time to develop post-traumatic ankle OA after a ligament sprains is 34 years [78]. This means that post-traumatic ankle OA affects relatively young individuals in their fourth or fifth decade of life. Ankle OA is associated with considerable pain and disability [79], particularly post-traumatic ankle OA [75]. This is particular concerning as post-traumatic ankle OA affects younger individuals more than primary ankle OA [75]. Compared to people without ankle problems, individuals with ankle OA report lower physical and mental health, and higher pain and disability [80]. Function in activities of daily living and sport/reactional activities, overall quality of life and independent living are all impaired in individuals with radiographic ankle OA compared to asymptomatic individuals [79]. Together, this research shows that ankle OA is a serious problem that greatly affects wellbeing.

34.4.2 Clinical findings There are no clinical diagnostic criteria for ankle OA. Clinically, ankle OA is usually diagnosed using a thorough assessment of the foot and ankle to determine that pain is originating from the ankle joint and to exclude the presence

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of pain solely in other structures (such as the tibialis posterior, Achilles, or peroneal tendons). Pain is commonly reported and reproduced by walking on flat, uneven ground or slopes, and ascending or descending stairs. Swelling around the ankle joint may be present and pain may be reproduced during palpation of the joint lines. Crepitus and/or pain may also occur during dorsiflexion and plantarflexion movements, particularly at end of range. A 2018 systematic review investigated physical impairments that characterize ankle OA [81]. Meta-analyses concluded that individuals with ankle OA present with considerably less sagittal plane (dorsiflexion/plantarflexion) range of motion and muscle strength of their affected ankle compared to their unaffected side and to individuals without ankle OA. Small decreases in calf circumference were also identified in individuals with ankle OA. In addition to deficits in ankle dorsiflexion and plantarflexion muscle strength and range of motion, other studies have identified deficits in heel raise (plantarflexion) endurance, inversion and eversion muscle strength [79,82], and range of motion [83]. While impairments in muscle function are multi-factorial, smaller cross-sectional area and greater fatty infiltration of the calf muscles in individuals with ankle OA may help to explain these strength and endurance deficits [84]. Impairments present in ankle OA negatively impact on ambulatory ability. Individuals with symptomatic ankle OA are slower to perform ambulatory tasks, such as walking 10 m on a flat surface and ascending and descending one flight of stairs, than those without ankle symptoms [79]. These functional deficits are consistent with greater self-reported disability and functional limitations performing activities of daily living and recreation/sport in individuals with ankle OA [79].

34.4.3 Structural and biomechanical features A recent systematic review has suggested that ankle OA is associated with a varus hindfoot position [81]. However, it is not known whether this feature is a consequence of ankle OA, or contributes to the development of the disease. A longitudinal study measured ankle joint alignment in individuals who had sustained a tibial fracture in the previous 6 12 years [85]. The authors identified that greater joint inclinations were associated with greater joint degeneration, higher levels of pain and greater restrictions in ankle joint range of motion, particularly for those with varus deformities. Malalignment at the ankle is thought to alter the distribution of load across the joint, and thus contribute to symptoms and impairments. A cadaver study showed that when the normal anatomy and alignment of the ankle is changed, the talocrural joint contact area decreases and the contact pressure increases [86]. Alternatively, it is well established that a cavovarus foot type is associated with lateral ankle injuries, the leading contributor to the high incidence of posttraumatic ankle OA [87]. Thus, associations between a varus hindfoot/cavovarus foot may simply be due to the confounding influence of a previous lateral ankle injury. Spatiotemporal, kinematic and kinetic alterations in gait have been identified in people with ankle OA. Individuals with ankle OA walk at a slower cadence, and have a reduced stride length and walking speed than those without ankle OA [88,89]. These spatiotemporal observations are supported by clinical findings of slower walking speed [79]. Several studies have shown that ankle OA is associated with reduced foot and ankle motion during gait. For instance, individuals with ankle OA have reduced total sagittal plane motion of the hindfoot, forefoot, and midfoot, and reduced heel strike and toe off pitch angles than those without ankle OA [88]. Research also suggests that people with ankle OA maintain the foot and ankle in a more dorsiflexed (less plantarflexed) position compared to controls [89,90]. In addition to reduced sagittal plane motion, there is evidence of reduced ankle internal/external rotation and inversion range of motion during gait in individuals with ankle OA [88,89,91]. These findings are consistent with clinical findings of deceased ankle range of motion [81,83] and reproduction of pain at end-range movement in ankle OA. Studies on gait kinetics indicate that peak anterior-posterior and vertical ground reaction forces, and the peak ankle dorsiflexion moment, are lower in ankle OA compared to healthy individuals [91]. This may be related to decreased ankle muscle strength [79,81,82] or a strategy to control pain during walking. Comparisons between sides in individuals with unilateral ankle OA suggests an asymmetrical gait pattern. The affected side has a shorter stance time, less sagittal and coronal plane motion, and less ankle plantarflexion at toe-off than the unaffected side [88,91]. Differences in stance time between limbs, and reduced anterior-posterior and vertical ground reaction forces, is suggestive of a limp, which may be a compensation for restricted movement of the affected side and/or a strategy to de-load and protect the affected ankle.

34.4.4 Clinical and biomechanical effects of conservative treatment There are no clinical practice guidelines, nor is there any evidence from high-quality randomized controlled trials to guide ankle OA management. A 2015 systematic review of the literature on non-surgical management of ankle OA did not identify any evidence-based interventions [92]. This review identified six RCTs, all which investigated the use of hyaluronic

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FIGURE 34.7 Three different types of custom-made foot orthoses used in the Huang et al. study (from left to right): an ankle foot orthosis, a rigid hindfoot orthosis, and an articulated hindfoot orthosis. From Huang, Y.-C., et al., Effects of ankle-foot orthoses on ankle and foot kinematics in patient with ankle osteoarthritis. Arch Phys Med Rehab, 2006. 87(5): p. 710 716 [96].

acid in ankle OA management. Analyses of data concluded that the quality of evidence was poor and there were insufficient data as to whether hyaluronic acid is beneficial in the management of ankle OA. Although clinical OA guidelines recommend education on the condition and self-management, physical activity and exercise, and weight loss if the person is overweight or obese [5,93], no studies have investigated the efficacy of these approaches for ankle OA. Of the six RCTs that investigated the use of hyaluronic acid in ankle OA, two also evaluated exercise management of ankle OA either alone or in combination. A small Turkish study (n 5 15 per group) compared hyaluronic acid injections to a home exercise program [94]. The exercise program consisted of isometric ankle dorsiflexion, plantarflexion, inversion and eversion, active ankle range of motion, stretching, and undescribed proprioceptive and closed kinetic chain exercises. They found that pain and physical activity improved similarly in both the hyaluronic acid and home exercise group, but walking distance was improved only in the exercise group [94]. A subsequent 2014 study compared a combined intervention of intraarticular hyaluronic acid and exercise to an intraarticular injection of Botulinum toxin type A [95]. The exercise program consisted of ankle range of motion, stretching, isometric and proprioceptive exercise. The study did not find any differences in pain or function between groups. While it may be speculated that foot/ankle orthoses could correct any malalignment at the foot and ankle and improve contact area and distribution of forces at the talocrural and subtalar joints, little is known about the effect of orthoses on symptoms and impairments in ankle OA. One study investigated the effect of three types of orthoses (custom-made ankle-foot orthoses, rigid hindfoot orthoses and articulated hindfoot orthoses) on gait kinematics in ankle OA (Fig. 34.7) [97]. The custom-made orthoses restricted sagittal plane motion (but not frontal or transverse plane motion) of the hindfoot and forefoot compared to a standard shoe when walking on flat ground. The other orthoses did not change hindfoot and forefoot motion in any movement plane. Researchers suggest that custom-made orthoses may be useful to restrict ankle motion during gait. However, the impact of restricting sagittal plane foot and ankle motion on pain and function, or kinematics of other joints of the lower limb, is not known.

34.5

Areas of future biomechanical research

Throughout this chapter, it has been highlighted that research on OA of the first MTPJ, midfoot, and ankle is substantially lacking, particularly when compared to OA at more proximal sites. This suggests the field has ample opportunities for future research, and perhaps some of the greatest opportunities for foot and/or ankle OA research lie in the field of biomechanics. These associations between biomechanics, joint changes, and OA symptoms highlight at least two promising areas for future foot/ankle OA biomechanical research. Firstly, a detailed understanding of intraarticular kinematics and in vivo contact forces and stress are needed. At the knee, a higher external knee adduction moment (a surrogate measure of in vivo medial knee load) has been associated with the development of knee pain in older people [98], with increased knee pain severity [99,100] and physical dysfunction in people with knee OA [101], and with greater odds of structural

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FIGURE 34.8 Finite element model of the foot and ankle musculoskeletal complex, including 30 bones, 85 ligament bundles with 1814 line elements, 74 cartilage layers, plantar fascia and encapsulated soft tissue (transparent). From Akrami, M., et al., Subject-specific finite element modeling of the human foot complex during walking: sensitivity analysis of material properties, boundary and loading conditions. Biomech Model Mechanobiol, 2018. 17(2): p. 559 576 [105].

OA progression [100,102,103]. However, there is now a growing recognition of the limitations of external moments to represent internal forces, and instead, researchers are increasingly using computational techniques such as musculoskeletal modeling to more accurately estimate in vivo tibiofemoral contact forces [104]. While these techniques are more problematic in the foot than the knee, largely due to the numerous articulations, very small joint moment arms and difficulties measuring intrinsic muscle activity (all used to inform the musculoskeletal models, amongst other variables), there are some promising avenues of research. For instance, a recent novel study has utilized subject-specific musculoskeletal modeling to calculate intrinsic foot muscle forces (derived from multi-segment foot kinematics and electromyography), combined with MRI-derived finite element modeling, to predict in vivo foot stresses and strain (Fig. 34.8) [105]. Other promising research has used high-speed biplanar fluoroscopy to accurately measure three-dimensional midtarsal joint kinematics, bone movement and surface velocity [106], and plantar aponeurosis elongation [107], in healthy populations. These, and other exciting biomechanical studies, offer new lines of research to help with our understating of the pathomechanics associated with OA of the foot and ankle. With this greater understanding of the pathomechanics of foot/ankle OA, new biomechanical interventions may be designed, and existing treatments optimized, to help relieve symptoms and improve function. The other promising area for future foot/ankle OA research is the need for longitudinal studies to help clarify whether biomechanical mechanisms cause or are a consequence of the disease. While many biomechanical features have been identified as common in people with first MTPJ, midfoot, and ankle OA, the vast majority of studies are cross sectional. Thus, the directional nature of associations between biomechanics and OA is unclear. Large longitudinal studies will help us understand whether commonly observed pathomechanical features lead to structural joint damage, OA pain, and other symptoms if they occur as adaptive mechanisms to reduce joints stress and symptoms, or if they are due to a third confounding factor. A good example of the latter point is the varus hindfoot posture observed in ankle OA. While this has been linked to ankle joint degeneration and pain, a cavovarus foot is also common in people who have experienced a lateral ankle injury [87]. Thus, associations between a varus hindfoot/cavovarus foot may instead be due to the pro-inflammatory mediators that are released following a previous lateral ankle injury [108], rather than longer-term increases in joint contact forces due to segmental biomechanics. This highlights that understanding temporal associations between biomechanics, joint changes, and symptoms will be important not only for preventing foot/ankle OA onset, but also to help slow disease progression, reduce symptoms and improve function.

34.6

Summary

This chapter has identified several biomechanical factors that have been associated with first MTPJ, midfoot, and/or ankle OA. For instance, anatomical features such as a dorsiflexed first metatarsal and longer phalanges are suggested to reduce first metatarsal gliding and rotation, causing the proximal phalanx to “jam” on to the first metatarsal head and increase joint compressive forces and structural change, which ultimately leads to pain in people with first MTPJ OA. In midfoot OA, a more everted foot posture and increased midfoot loading are speculated to increase dorsal midfoot compressive forces, also leading to joint changes and pain. In contrast, a more varus hindfoot has been identified in

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people with ankle OA, and this is suggested to increase talocrural joint contact pressures and consequently to cause ankle OA symptoms and associated disability.

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[40] Pons M, et al. Sodium hyaluronate in the treatment of hallux rigidus. A single-blind, randomized study. Foot Ankle Int 2007;28(1):38 42. [41] Paterson KL et al. Podiatry intervention vs usual general practitioner care for symptomatic radiographic osteoarthritis of the first metatarsophalangeal joint: a randomised clinical feasibility study. Arthritis Care Res (Hoboken), 2019. Accepted Author Manuscript. [42] Munteanu SE, et al. Shoe-stiffening inserts for first metatarsophalangeal joint osteoarthritis (the SIMPLE trial): study protocol for a randomised controlled trial. Trials 2017;18(1):198. [43] Menz HB, et al. Rocker-sole footwear vs prefabricated foot orthoses for the treatment of pain associated with first metatarsophalangeal joint osteoarthritis: study protocol for a randomised trial. BMC Musculoskelet Disord 2014;15(1):86. [44] Hutchins S, et al. The biomechanics and clinical efficacy of footwear adapted with rocker profiles—evidence in the literature. 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[64] Kitaoka HB, Luo ZP, Kai-Nan. Contact features of the talonavicular joint of the foot. Clin Orthop Relat Res 1996;325:290 5. [65] Rao S, Baumhauer JF, Nawoczenski DA. Is barefoot regional plantar loading related to self-reported foot pain in patients with midfoot osteoarthritis. Osteoarthr Cartil 2011;19(8):1019 25. [66] Yi T, et al. Effect of full-length carbon fiber insoles on lower limb kinetics in patients with midfoot osteoarthritis: a pilot study. Am J Phys Med Rehab 2018;97(3):192 9. [67] Arnold JB, et al. Foot and leg muscle weakness in individuals with midfoot pain and osteoarthritis. Osteoarthr Cartil 2019;27:S449. [68] Paterson KL, et al. Management of foot/ankle osteoarthritis by Australian general practitioners: an analysis of national patient-encounter records. Osteoarthr Cartil 2018;26(7):888 94. [69] Rao S, et al. Orthoses alter in vivo segmental foot kinematics during walking in patients with midfoot arthritis. Arch Phys Med Rehab 2010; 91(4):608 14. 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Chapter 35

Diabetic Foot Disease Bijan Najafi1 and Gu Eon Kang1,2 1

Interdisciplinary Consortium on Advanced Motion Performance (iCAMP), Division of Vascular Surgery and Endovascular Therapy, Michael

E. DeBakey Department of Surgery, Baylor College of Medicine, Houston, TX, United States, 2Department of Bioengineering, University of Texas, Dallas, TX, United States

Abstract Diabetes is a major health concern with a considerable and growing impact worldwide. Diabetic foot disease is one of the most serious complications that can arise from untreated or poorly controlled diabetes, and it is associated with increases in morbidity and mortality. It can involve loss of sensation in the feet, vascular disease, and can lead to deformities, plantar ulcers, and amputations. This chapter discusses the background, risk factors, and consequences of diabetic foot disease, with a particular focus on biomechanical factors including changes in gait and tissue properties. Future areas of research related to this area are also discussed.

35.1

Background on diabetes

Diabetes is a significant public health concern that places considerable burdens on our society. Over the last few decades, the number of individuals with diabetes has been increasing alarmingly. The Centers for Disease Control and Prevention estimated that in 2018 more than 10% of the US population ($ 34 million people) had diabetes and nearly 35% of the US population (B90 million people) was at high risk of developing diabetes (i.e., prediabetes) [1]. The total direct (i.e., medical cost) and indirect (i.e., reduced productivity) economic cost of diabetes is enormous, accounting for $327 billion annually in the US alone [2] and nearly $1.5 trillion annually worldwide [3]. Diabetes refers to a group of metabolic disorders characterized by high glucose levels in the blood (i.e., hyperglycemia) caused by deficits in insulin secretion and/or action [4]. The vast majority of diabetes falls into two categories based on pathogenetic mechanisms: type 1 and type 2 diabetes. Type 1 diabetes, formerly known as insulin-dependent diabetes, is characterized by immune destruction of β cells (insulin secreting cells) and is caused by an absolute lack of insulin secretion [5]. Type 2 diabetes, also known as non-insulin-dependent diabetes, is caused by a combination of abnormal patterns of insulin secretion and insulin resistance. Type 2 diabetes is much more common than type 1 diabetes, and it accounts for 90% 95% of all cases of diabetes [6]. The diagnosis of diabetes is based on the degree of hyperglycemia and relies on measuring glucose levels in the fasting state (i.e., no caloric intake for minimum 8 hours) or from an oral glucose tolerance test (i.e., a metabolic stress test) [7]. For example, a diagnosis of diabetes can be made with a fasting glucose level $ 126 mg/dL (a fasting glucose level of 110 125 mg/dL for prediabetes) or with a glucose level $ 200 mg/dL from an oral glucose tolerance test (140 199 mg/dL for prediabetes) [7]. A diagnosis of diabetes can also be based on chronic glycemia, often measured with a glycated hemoglobin A1C levels (i.e., average glucose levels in the past 8 to 12 weeks; hemoglobin A1C $ 6.5% for diabetes and 5.7% 6.4% for prediabetes) [8]. Several risk factors that increase the chance of developing prediabetes and, ultimately, diabetes, particularly type 2 diabetes, have been identified. Although there exist non-modifiable risk factors such as familial history of diabetes, aging, or racial or ethnic background [9], some of these factors are related to an individual’s modifiable lifestyle. For example, numerous studies demonstrated overweight/obesity is the top modifiable risk factor for diabetes [10], and several national and international institutions, including the National Institute of Diabetes and Digestive and Kidney Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00025-1 © 2023 Elsevier Inc. All rights reserved.

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Exercise, proper footwear, and regular care are key to prevent diabetic plantar ulcer

About 422 million people worldwide have diabetes

Loss of “the gift of pain” due to neuropathy is a major cause of diabetic plantar ulcer

At least 70% of amputations are potentially preventable

Diabetic planter ulcer may also put patients at risk for other adverse events such as falls

1 in 3 people with diabetes have the risk of diabetic plantar ulcer

Diabetic plantar ulcers underlie 85% of all lower extremity amputations

Rate of recurrence : 65% in 5 years

FIGURE 35.1 Diabetic foot infographic.

Diseases, provide clinical guidelines of healthy body mass index that reduce the risk of developing diabetes [11]. Along with overweight/obesity, physical inactivity (i.e., sedentary lifestyle) is another top risk factor for diabetes [12]. The American Heart Association and American Diabetes Association recommend a minimum of 150 minutes of weekly aerobic exercise with moderate intensity or a minimum of 75 minutes of weekly aerobic exercise with vigorous intensity to reduce risk for diabetes [13]. An unhealthy diet is also positioned amongst the top risk factors. For example, research has shown that low cholesterol absorption may contribute to the insulin resistance and is associated with increased risk for diabetes [14]. Similar to weight and physical activity management, dietary control of cholesterol is recommended as a target for diabetes management [15]. If left untreated or poorly managed, individuals with diabetes can develop serious complications (Fig. 35.1). Diabetic foot syndrome, defined as “infection, ulceration, or destruction of tissues of the foot of an individual with diabetes” by the International Working Group on the Diabetic Foot [16], is one of the most devastating diabetic complications that significantly reduces quality of life and increases the morbidity and mortality of the affected individuals. This chapter provides an overview of diabetic foot disease, biomechanical abnormalities caused by diabetic foot disease, and directions for future biomechanics research to advance our current understanding of the management of diabetic foot disease.

35.2

Overview of key negative outcomes of diabetic foot disease

35.2.1 Diabetic peripheral neuropathy Diabetic peripheral neuropathy is one of the most common complications of diabetes [17], and it affects up to half of all individuals with diabetes [18]. Diabetic peripheral neuropathy is peripheral nerve dysfunction from diabetes [16]. Symptoms of diabetic peripheral neuropathy may be painful (e.g., burning pain, electrical sensation) and non-painful (e.g., numbness, tingling) [18]. Diabetic peripheral neuropathy is a leading risk factor for subsequent plantar ulcers and their recurrence, accounting for up to 60% of all ulcers [19,20]. Studies have also shown that individuals with diabetic peripheral neuropathy may also experience absent ankle reflexes and loss of muscle strength [21].

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Diabetic peripheral neuropathy results in a loss of sensorimotor function, which is manifested in reduced plantar sensation [22,23] and deteriorated static and dynamic balance control during locomotor tasks [24 27], in turn increasing the risk of fall in the affected individuals [28]. For example, individuals with diabetic peripheral neuropathy are 23 times more likely to fall than age-matched normal individuals [29]. Furthermore, it has been reported that the reduced plantar sensation is an independent factor to increase the risk of falls in older adults with diabetes [30]. The reduced sensorimotor function and increased risk of falls associated with diabetic peripheral neuropathy triggers the development of fear of falling, and further reduces physical activity [31 33]. This vicious circle among diabetic peripheral neuropathy, loss of sensorimotor function, fear of falling, and reduced physical activity further worsens and progresses symptoms of diabetic peripheral neuropathy and places more burdens and pressures on the affected individuals, their significant ones, care providers, and eventually health care systems [34,35].

35.2.2 Peripheral vascular disease Individuals with diabetes are at higher risk of developing peripheral vascular disease, which is known as a multifactorial cause for developing diabetic foot disease. Peripheral vascular disease is a condition characterized by atherosclerotic occlusive disease in the leg and feet and affects more than 8 10 million adults in the US [36]. At least 20% of individuals with peripheral vascular disease have diabetes [37]. Studies have shown that nearly 50% of individuals with a diabetic plantar ulcer have peripheral arterial disease [38]. For individuals with severe peripheral vascular disease, a significant blockage in the lower extremity blood vessels may cause pain during resting states and cause gangrene of the lower extremities, called critical limb ischemia [17], affecting more than 10% of individuals with peripheral arterial disease [39]. High lower extremity amputation rates are a major negative outcome caused by critical limb ischemia, the end stage of peripheral arterial disease: up to 38% within 1 year [40]. It is estimated that mortality rates after lower extremity amputation due to peripheral arterial disease increase to nearly 30% for acute mortality and more than 70% for 5-year mortality [41]. The most classic manifestation of peripheral vascular disease is intermittent claudication that is characterized by pain, cramping, or fatigue in the lower extremity. Owing to the pain in the lower extremity, intermittent claudication is associated with functional limitations, often measured by reduced pain-free walking distance [42] or 6-minute walk distance [42], or impaired dynamic and static balance in some cases [43,44].

35.2.3 Diabetic plantar ulceration Diabetic plantar ulcers, which are open sores or lesions on the foot that do not heal for a long period of time, 15 20 weeks on average [45], are the most frequently recognized lower extremity complication of diabetes [19,20]. It is estimated that up to 34% of individuals with diabetes will develop diabetic plantar ulcers during their lifetime [20]. Diabetic plantar ulceration is one of the most devastating complications of diabetes, accounting for up to 70% of all lower extremity amputations [20]. It is also estimated that diabetic plantar ulcers increase the risk of death within 5 years up to 2.5 times compared to individuals with diabetes who do not have a plantar ulcer [20]. Another devastating characteristic of a diabetic plantar ulcer is high recurrence rates. Nearly 40% of individuals who have a healed plantar ulcer will develop a recurrent plantar ulcer within 1 year, and this recurrence rate increases to 60 and 65% within 3 years and 5 years, respectively [20]. The vast majority of diabetic plantar ulcers are caused by diabetic peripheral neuropathy (alone or together with foot ischemia). The most common pathway of a diabetic plantar ulcer is sensorimotor dysfunction that leads to high plantar pressure, which in turn results in foot deformities and gait impairment (Fig. 35.2) [46,47]. The foot deformities and gait impairment will then potentially further heighten plantar pressures. This vicious cycle contributes to the development, progression, and recurrence of diabetic plantar ulcers.

FIGURE 35.2 Diabetic plantar ulcers rely on repetitive loading over an area that is subject to high vertical or shear stress.

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35.2.4 Foot deformities Foot deformities such as flat foot, hammertoes, and hallux valgus are common in individuals with diabetes. It is estimated that foot deformities together with loss of sensorimotor function increases the risk of lower extremity amputation by up to 17 times [48]. The pathogenesis of foot deformities in individuals with diabetes is not well understood; however, it is commonly accepted that joint immobility, which causes thickening of the Achilles tendon and plantar fascia, and muscle atrophy and weakness caused by peripheral neuropathy may result in foot deformities [48]. Owing to the structural alterations in the foot and ankle, foot deformities may alter gait patterns independently from diabetic peripheral neuropathy. For example, foot deformities, which are associated with thickened Achilles tendon and plantar fascia in particular, may cause lesser adaptability on walking surface, manifested as a prolonged stance phase during gait, and result in an unsteady gait pattern [48].

35.2.5 Lower-extremity fractures Diabetes may increase the risk of fracture in the lower extremity. For example, several large cohort studies found that women with diabetes had heightened risk of hip and/or foot fracture compared to age-matched nondiabetic older women [49,50]. In addition, a large case-control study found that individuals with diabetes (both men and women) had 40% 70% greater risk of hip fracture compared to age- and sex-matched nondiabetic individuals [51]. However, there is very little published information concerning the deleterious effect of diabetes on fracture healing or its overall effects on the structural integrity of articular cartilage and bone. There are very few clinical reports that address fracture complications in diabetes and no prospective work that reports long term clinical outcomes. Most of the clinical reports in humans with diabetes and fracture complications focus on ankle fractures and small case series of patients with Charcot neuroarthropathy of the foot and ankle in patients with type 2 diabetes [52,53]. In a retrospective study, Liu et al. [54] studied 17,464 patients undergoing surgery for ankle fractures. Of these patients, 2044 (11.7%) had diabetes and 15,420 (88.3%) did not. They excluded patients older than 90 years or with inadequate perioperative data. Their results revealed very high rates of non-unions, wounds, infections, joint pain, and the need for long-term bracing in patients with diabetes compared to those without. For instance, the risk of non-union was 2.8 times higher (confidence interval: 2.11 3.37) in those with diabetes compared to those without. Likewise, in a retrospective study, Glassman compared complications after lumbar spinal fusions in adults with type 2 diabetes and adults without diabetes. The incidence of non-union was 4 5 times higher in patients with diabetes [55]. The literature regarding the mechanism of action of diabetes on the musculoskeletal system is severely deficient. Even though the vast majority of fracture complications are in persons with type 2 diabetes, most of the animal studies are in type 1 models [56,57]. In this context, type 2 diabetes has not been well studied. The metabolic abnormalities and response to fracture may be significantly different in type 1 and type 2 diabetes. Therefore, the existing work in type 1 diabetic models should be validated in type 2 diabetic animal models. This should improve the translational benefit to human evaluation and interventions. To our knowledge none of the published studies in animals with type 1 diabetes evaluate the effect of peripheral neuropathy or peripheral arterial disease on bone repair. The existing studies evaluate young animals 2 or 3 weeks after diabetes has been induced and before peripheral neuropathy, or microvascular and macrovascular disease develop. The presence and severity of neuropathy and vascular disease, in conjunction with diabetes, have been shown to increase the risk of wounds, infections, and amputations in clinical practice, when compared to the presence of diabetes alone [58,59]. We anticipate the same would be true of fracture complications; however, clinical studies in this area are scarce.

35.2.6 Charcot neuroarthropathy Charcot neuroarthropathy, also known as Charcot foot, is characterized by the earliest phase of inflammation affecting the bones, joints, and soft tissues in the foot and ankle (Fig. 35.3) [60]. Today, diabetes and neuropathy is the commonest cause of Charcot foot in western countries and frequently leads to foot ulceration and amputation [61 63]. This condition can have a major impact on activities of daily living and quality of life, and its presence has been associated with significant mortality [62]. The pathogenesis is not well understood, but it is generally agreed trauma (micro-repetitive or single-instance macrotrauma) in the presence of neuropathy is required to spark this syndrome [64]. Trauma precipitates an inflammatory cascade and vasodilation [65]. The vasodilation may “wash out” minerals from the bone reducing density and thereby making it more susceptible to fractures. This, coupled with glycosylation of periarticular

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Diabetic neuropathy “Resultant hyperemia”

Inflammaon

Exaggeration of blood flow

Continued Insult

CHARCOT FOOT

Minor trauma Foot ulcer Surgical insult or revascularization

Fracture / Dislocation FIGURE 35.3 Charcot foot is an inflammatory syndrome characterized by varying degrees of bone and joint disorganization secondary to underlying neuropathy, trauma, and perturbations of bone metabolism.

tissue, particularly the Achilles tendon, may make the person with Charcot arthropathy of the foot prone to its characteristic fracture-dislocations. The hallmark feature of Charcot foot is midfoot collapse, “rocker bottom foot,” accompanied by pain and discomfort [60]. The incidence rate of Charcot foot is approximately 1% among individuals with diabetes [66]. It has been estimated that more than 60% of individuals with diabetes who have Charcot foot will develop a plantar ulcer [67]. If an individual has both Charcot foot and a plantar ulcer, the risk of lower extremity amputation increases by 12 times compared to those who have Charcot foot alone [68]. Treatment of Charcot foot depends on the stage during which it is diagnosed [63]. Early diagnosis is challenging but allows the opportunity to intervene and prevent development or worsening of deformities. Several studies suggested the presence of unilateral heat and swelling in a neuropathic diabetic patient should be presumed to be due to acute Charcot neuropathy until proven otherwise [63,69,70]. More specifically, in acute Charcot neuropathic patients, the foot is at least 2 C (approximately 4 F) hotter than the contralateral foot. For example, Armstrong and colleagues [70] used a portable infrared skin temperature device to demonstrate a mean difference of 8.3 F (B4.6 C) between the affected foot and contralateral foot in 21 Charcot neuropathic patients. This difference was 5.6 F (B3.1 C) in patients with neuropathic ulcers (n 5 44), and zero in patients with asymptomatic sensory neuropathy (n 5 76). The presence of unilateral heat in Charcot neuropathic patients was explained by local repetitive stress cause by joint abnormality on the affected side [70]. Najafi et al. [71] demonstrated that accuracy of acute Charcot neuropathy detection could be increased by assessing plantar temperature response to walking, called the thermal stress response test. In summary, they examined dynamic changes in plantar temperature as a function of graduated walking activity to quantify thermal stress responses during the first 200 steps. Fifteen individuals with Charcot neuropathy and 17 non-Charcot neuropathic participants with type 2 diabetes and peripheral neuropathy were recruited. All participants walked for two predefined paths of 50 and 150 steps. A thermal image was acquired at baseline, after acclimatization, and immediately after each walking trial. The hot spot temperature was identified by the 95th percentile of measured temperature at each anatomical region (hind/mid/forefoot). They reported that during initial activity, the thermal stress response is reduced in all participants, but the temperature drop for the non-affected foot was 1.9 times greater than the affected side in the Charcot neuropathic group. Interestingly, the thermal stress response in Charcot neuropathy was sharply increased after 50 steps for both feet, while no difference was observed in non-Charcot neuropathy between 50 and 200 steps. They concluded that the plantar stress response may be helpful to differentiate Charcot neuropathy and its response to treatment earlier in its course. Later, the same team suggested that thermal stress response could be monitored during activities of daily living using smart socks that enable the simultaneous measurement of pressure and temperature [72]. Several commercially available technologies could also enable remote assessment of plantar temperature during activities of daily living, thus improving implementation of this concept for early detection of Charcot foot [73]. However, this hypothesis needs to be confirmed in future clinical studies.

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The distribution and magnitude of plantar pressure during gait can also provide insight into the functional manifestations of foot and ankle disorders and may be used for the early diagnosis of abnormal foot biomechanics due to Charcot foot. Additionally, these dynamic measurements often serve as objective measures for outcome evaluation following foot surgery and can be used to track disease progression [74 78]. However, plantar pressure magnitude alone may be inappropriate when studying the effects of foot surgery. Following corrective foot surgery, patients will often increase their gait speed as a result of greater confidence, stability, and a more efficient gait pattern. Although this increase may be practically advantageous, it may also increase plantar pressure magnitude, historically viewed as a negative outcome. For a fair evaluation of the success of surgery, a novel and speed independent variable is required that reflects improved foot biomechanics post-surgery. To address limitation of plantar pressure magnitude for monitoring abnormal plantar pressure because of Charcot Foot, Najafi et al. [79] introduced a dynamic plantar loading index, which represents the similarity of the actual pressure distribution with a normal distribution. The dynamic plantar loading index may range from negative 1 to positive 1 and as the value increases positively so does the similarity between the actual and normalized pressure distributions. They tested this novel score on the plantar pressure pattern of healthy subjects (n 5 15), Charcot patients pre operation (n 5 4), and a Charcot patient post foot reconstruction (n 5 1). In healthy subjects, the dynamic plantar loading index was 0.46 6 0.1. When subjects increased their gait speed by 29%, plantar pressure magnitude was increased by 8%, while the dynamic plantar loading index was not changed, suggesting that the dynamic plantar loading index is independent of gait speed. In preoperative Charcot patients, the dynamic plantar loading index was , 0, however, the dynamic plantar loading index increased post-surgery (dynamic plantar loading index 5 0.42), indicating a transition to a more normal plantar pressure distribution after Charcot reconstruction. In follow-up studies [80,81], they demonstrated that the dynamic plantar loading index is independent of gait speed, is associated with severity of foot deformity, and could better describe pain relief in response to custom made orthoses compared to conventional plantar pressure magnitude measures such as peak pressure and pressure time integral. They also speculated that abnormal pressure distribution in Charcot neuropathic patients may explain the unilateral increase in plantar temperature. With advances in wearable technologies such as smart socks and smart insoles [73], enabling continuous monitoring plantar temperature and pressure, we anticipate these metrics could be utilized to better identify early signs of foot deformity including Charcot foot.

35.2.7 Lower extremity amputation More than 100,000 lower extremity amputations are performed every year in the US [82]. Of the amputations performed, diabetes is the most frequent cause of non-traumatic lower extremity amputations, accounting for up to 90% of all amputations [38]. The major cause of the high amputation risk is thought to be a combination of peripheral neuropathy, infections, and non-healing plantar ulcers stemming from diabetes (Fig. 35.4) [20]. The in-hospital and 30-day

FIGURE 35.4 Contributing factors to lower extremity amputation.

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mortality after lower extremity amputation in individuals with diabetes is greater than in individuals who received a cardiac surgery (coronary artery bypass graft surgery) [82]. The 1-year mortality in individuals who underwent major lower extremity amputations (either below or above knee) was more than 40% [38].

35.3

Risk factors for the development and progression of diabetic foot disease

Although the spectrum of diabetic foot disease varies and causes of the development and progression of diabetic foot disease are multifactorial, diabetic peripheral neuropathy and peripheral vascular disease are generally considered as the major risk factors for diabetic plantar ulcers (Fig. 35.4) [16,83]. Diabetic peripheral neuropathy causes foot deformities and abnormal biomechanical loading patterns on the foot. With the presence of diabetic peripheral neuropathy, even a minor injury such as a thermal injury from poorly fitting shoes can lead to an ulcer. Because individuals with diabetic peripheral neuropathy have usually lost the ability to feel pain, also described as “the gift of pain” by Dr. Paul Brand [84], awareness of local inflammation or injury on the foot is delayed and the time interval between an injury and an intervention (either surgical or non-surgical) becomes much longer in individuals with diabetic peripheral neuropathy compared to those with no neuropathic symptoms [83]. The lack of pain sensation means unchecked repetitive plantar pressures are applied to foot injuries in individuals with diabetic peripheral neuropathy [19,46], which then significantly contribute to the progression of plantar ulcers. Furthermore, diabetic peripheral neuropathy is usually not resolved even after a plantar ulcer is healed, and thus these patterns continue to happen, and, consequently, individuals with diabetic peripheral neuropathy are prone to the recurrence of plantar ulcers [20]. Peripheral vascular disease is another independent risk factor from diabetic peripheral neuropathy that heightens the risk of diabetic plantar ulcers and subsequent lower extremity amputation [20], although peripheral vascular disease rarely causes plantar ulcers by itself [85]. When peripheral vascular disease is combined with minor trauma, the combination commonly leads to hard-to-heal ulcers [85]. Minor trauma to the foot or ankle and possible subsequent infection would increase the demand on oxygen supply [85]. However, with the presence of peripheral vascular disease, this demand becomes beyond the circulatory capacity of the affected individuals, which then leads to plantar ulcerations and ultimately to lower extremity amputation [85]. Like diabetic peripheral neuropathy, peripheral vascular disease is generally not resolved after ulcer healing and contributes to ulcer recurrence [20].

35.4

Changes in kinematics and kinetics in diabetic foot disease

While walking, the foot absorbs or dissipates impact forces from contact with the walking surface and helps our body to adapt to the walking surface [86]. Normal and smooth articulations among the subtalar, talocrural, and talocalcaneonavicular joints play the major function of controlling the relative movement of the lower leg to the foot in the sagittal plane (i.e., ankle dorsiflexion-plantarflexion) [87,88]. These normal and smooth articulations in the foot and ankle are disrupted in individuals with diabetic foot disease, leading to impairments in function (Fig. 35.5) [47,87,89]. Numerous studies have reported gait impairment in individuals with diabetic peripheral neuropathy. For example, Menz and colleagues compared gait parameters between a group with diabetic peripheral neuropathy and a group of age-matched controls on level and uneven surfaces [90]. The authors found significantly slower gait speed and significantly reduced rhythmic gait pattern for the diabetic peripheral neuropathy group compared to the control group on both surfaces. In another study, Richardson and colleagues investigated gait stability and variability in older women with diabetic peripheral neuropathy and compared that to those without diabetic peripheral neuropathy [91]. They found unstable and more variable gait patterns (i.e., greater step width-tostep length ratio) in older women with diabetic peripheral neuropathy compared to those without diabetic peripheral neuropathy. Additionally, Najafi and colleagues compared barefoot walking over 20 meters between individuals with diabetic peripheral neuropathy and healthy controls [92]. They found that individuals with diabetic peripheral neuropathy had increased gait speed variability, and this variability had a strong correlation with the severity of peripheral neuropathy, as measured by vibration perception threshold. More recently, Kang and colleagues investigated the gait initiation phase of older adults with diabetic peripheral neuropathy ($ 65 years old) and compared them to agematched normal controls [24]. They found that older adults with diabetic peripheral neuropathy took significantly more steps and a longer distance to get into steady-state gait phase and had a more unstable gait initiation phase (i.e., larger medio-lateral sway) than normal controls. These impairments in gait and balance in people with diabetic peripheral neuropathy may be associated with developing fear of falling. For example, Lalli and colleagues investigated the presence of fear of falling (based on a singleitem questionnaire: Do you have a fear of falling?) in people with diabetes (n 5 20), diabetic peripheral neuropathy

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FIGURE 35.5 Biomechanical dysfunction in diabetic foot disease.

without neuropathic pain (n 5 20), and diabetic peripheral neuropathy with neuropathic pain (n 5 22) [27]. They found that the presence of fear of falling was more than 12 times and 3 times that in people with diabetic peripheral neuropathy with neuropathic pain (64%) compared to those with diabetes (5%) and diabetic peripheral neuropathy without neuropathic pain (20%), respectively. Similar findings were reported by Kelly and colleagues [31]. In their study, the severity of fear of falling in 16 people with diabetic peripheral neuropathy was examined using the Falls Efficacy Scale-International [93]. Of those samples, 81% had moderate-to-high fear of falling. However, interestingly, they also found that the severity of fear of falling was not related to the level of peripheral neuropathy (i.e., vibration perception threshold in the plantar surface). As the presence of fear of falling further worsens motor performance [94], the level of physical activity seems to be further restricted. Kang and Najafi measured the level of physical activity objectively using a chest-worn activity monitor in people with neuropathy (n 5 49) and demonstrated the duration of walking activity and total step counts are less in people with neuropathy with high fear of falling (as measured by the Falls Efficacy Scale-International) compared to those with low fear of falling [33]. Furthermore, they found that total step counts are significantly correlated with the severity of fear of falling. In three-dimensional clinical gait analysis, the foot is usually designed as one rigid body segment. Although the one-segment foot model can provide the overall function during gait, given the anatomical complexity of a foot that is made up of 26 bones, it focuses more on the ankle movement rather than the function of the foot [95]. Multi-segment foot models overcome the limitations of one-segment foot models. In multi-segment foot models, the foot is often modeled using at least three segments excluding the tibia (e.g., hindfoot, forefoot, and hallux) [96], and a few studies have utilized multi-segment foot models for gait analysis in individuals with diabetes. Rao and colleagues segmented the foot into forefoot, the first metatarsal, and calcaneus segments, and compared foot kinematics between individuals with diabetes and neuropathy and age-, sex- and body mass index-matched nondiabetic controls while they were walking at a controlled speed (0.89 m/s) [97]. They found that individuals with diabetes and neuropathy walked with reduced frontal and sagittal plane ranges of motion of the calcaneus relative to the tibia, and reduced frontal and transverse ranges of motion of the first metatarsal relative to the calcaneus. In another study, Sawacha and colleagues investigated foot kinematics during gait in individuals with diabetic peripheral neuropathy and nondiabetic controls using a four-segment model, which was composed of the tibia, hindfoot, midfoot, and forefoot [98]. They found significant differences in forefoot and midfoot flexion between the two groups. Deschamps and colleagues utilized a three-segment foot model (calcaneus, midfoot, and metatarsus) and investigated sagittal, frontal, and transverse ranges of motion during gait in three groups: individuals with diabetes and neuropathy, individuals with diabetes without neuropathy, and nondiabetic individuals [99]. They found that individuals with diabetes and neuropathy had significantly reduced sagittal plane range of motion of the calcaneus segment during the forefoot push-off phase compared to both individuals with diabetes without neuropathy and nondiabetic individuals.

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Another major function of the foot during gait is weightbearing. In individuals with diabetic foot disease, due to loss of sensorimotor function, foot deformities, muscle weakness, and possible presence of plantar ulcers, the biomechanical loading pattern on the plantar surface is altered [47,87,89,100]. Since most plantar ulcers associated with diabetic peripheral neuropathy occur around the forefoot (the toes and metatarsal heads), abnormal loading patterns on this area has been of particular interest [89]. For example, Caselli and colleagues investigated the forefoot and hindfoot plantar pressure and forefoot-to-hindfoot plantar pressure ratio during gait in individuals with diabetes with no neuropathy, mild neuropathy, moderate neuropathy, and severe neuropathy [101]. They found that both forefoot and hindfoot plantar pressure increases with the severity of neuropathy. They also found that the forefoot-to-hindfoot plantar pressure ratio during gait was predictive of the development of plantar ulceration within a 30-month period. Furthermore, Sacco and colleagues examined plantar pressure during barefoot walking among four groups with diabetes with no neuropathy, mild neuropathy, moderate neuropathy, and severe neuropathy [102]. They found that the groups with moderate or severe neuropathy had greater peak pressure and pressure-time integral in the forefoot and heel compared to the groups with no neuropathy or mild neuropathy.

35.5

Changes in tissue characteristics

35.5.1 Muscle: fatty infiltration and reduction of intrinsic foot muscle volumes Several studies have reported severe weakness in the intrinsic muscles of the diabetic foot, which subsequently leads to foot deformities in individuals with diabetic peripheral neuropathy [103]. Foot deformities such as hammer toes and claw toes may result in the fat pads under the metatarsal heads moving distally, which contributes to the metatarsal heads only being covered by a thin layer of soft tissue and leads elevated plantar pressure during gait [89]. Suzuki and colleagues found structural changes in plantar muscles and greater fat content in individuals with diabetes and neuropathic ulcers compared to those without neuropathic ulcers [104]. Bus and colleagues reported severe atrophy in the intrinsic foot muscles of individuals with diabetic peripheral neuropathy [105]. In a subsequent study, Bus and colleagues reported thinner sub-metatarsal fat pads and thicker sub-phalangeal fat pads for individuals with diabetic peripheral neuropathy and toe deformities than for individuals with diabetic peripheral neuropathy and no toe deformities [106]. Andersen and colleagues measured the total volume of the intrinsic muscles of the foot in individuals with diabetic peripheral neuropathy, individuals with diabetes and no neuropathy, and nondiabetic individuals [107]. They found the total volume of the intrinsic foot muscles were nearly 50% less for individuals with diabetic peripheral neuropathy compared to the other groups.

35.5.2 Bone Diabetes (both type 1 and type 2) increases the risk of bone fracture, yet the underlying link between the condition and bone fracture risk is not fully understood [108]. Poor bone mineral density is often suspected as a risk factor of bone fracture, but in diabetes this is controversial. For example, bone mineral density decreases in individuals with type 1 diabetes, but increases or is preserved in individuals with type 2 diabetes [108]. The impact of peripheral neuropathy on bone mineral density is also unclear [109]. Rix and colleagues investigated bone mineral density in individuals with diabetic peripheral neuropathy (type 1 diabetes), individuals with type 1 diabetes and no-to-mild neuropathic symptoms, and nondiabetic controls [110]. They found lower bone mineral density in individuals with diabetic peripheral neuropathy compared to the other two groups. In contrast, Rasul and colleagues compared bone mineral density between individuals with type 2 diabetes with and without polyneuropathy and found similar bone mineral density between the two groups [111].

35.5.3 Cartilage Diabetes, particularly type 2 diabetes, is an independent risk factor for developing osteoarthritis [112]. One possible cause for this link is high glucose concentration and insulin resistance that induces oxidative stress and contributes to cartilage destruction [112]. However, evidence regarding diabetes, osteoarthritis, and cartilage destruction has focused on the hip and knee, and in the foot and ankle research is extremely sparse, which we will discuss later in future research areas.

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35.5.4 Tendon Several studies reported changes in tendon structures in individuals with diabetic foot disease. Bolton and colleagues investigated the length and thickness of the plantar aponeurosis and flexor hallucis longus tendon in individuals with diabetic peripheral neuropathy and age-, sex-, body mass index- and foot size-matched nondiabetic controls [113]. They found thicker plantar aponeurosis and flexor hallucis longus tendons in individuals with diabetic peripheral neuropathy compared to the controls. Giacomozzi and colleagues investigated the thickness of the Achilles tendon among individuals with diabetes and no neuropathy, individuals with diabetic peripheral neuropathy, individuals with diabetic peripheral neuropathy and a history of plantar ulcer, and nondiabetic healthy controls [114]. They found thicker Achilles tendons in individuals with diabetic peripheral neuropathy (both with and without a history of plantar ulcer) compared to the other two groups.

35.5.5 Plantar fascia The plantar fascia is a ligament in the foot, which has an important role in biomechanical loading during gait. Diabetes alters the structure of the plantar fascia. D’Ambrogi and colleagues investigated the thickness of the plantar fascia in individuals with diabetic peripheral neuropathy and a history of plantar ulcer, individuals with diabetic peripheral neuropathy, individuals with diabetes and no neuropathy, and healthy controls [115]. They found that the plantar fascia was thicker in individuals with diabetes (regardless of neuropathic symptoms) compared to healthy controls, but there was no difference among the groups with diabetes. These were consistent with results from other studies reported by Bolton and colleagues [113], and Giacomozzi and colleagues [114]. More recently, Kimura and colleagues investigated intrinsic foot muscle volume and the thickness of the plantar aponeuroses in individuals with diabetic peripheral neuropathy and claw toes, individuals with diabetic peripheral neuropathy without claw toes, non-neuropathic individuals with claw toes, and non-neuropathic individuals without claw toes [103]. They found that, among the four groups, individuals with diabetic peripheral neuropathy with claw toes had the smallest intrinsic muscle volume and the thickest plantar aponeuroses.

35.6

The relationship between foot deformities and plantar ulceration

Several studies have examined the indirect and direct relationship between foot deformities and plantar ulceration. In a prospective cohort study, Boyko and colleagues examined independent risk factors of plantar ulcers among individuals with diabetes and no plantar ulcer [116]. They identified that the Charcot foot had an independent association with ulcer development within 18 months. Ledoux and colleagues investigated the association between foot deformities and ulcer development in individuals with diabetes who were at high-risk of diabetic foot disease [117]. They found that foot deformities such as claw toes, hammer toes, and hallux limitus was associated with ulcer development within 2 years. These findings were confirmed in recent large cohort studies. Yazdanpanah and colleagues conducted a populationbased study aimed at identifying independent risk factors for diabetic plantar ulcer in individuals with diabetes, and reported foot deformities such as hammer toe, hallux valgus, bunion, prominent metatarsal head, and Charcot foot was associated with ulcer development within 2 years [118].

35.7 The relationship between lower extremity fractures and Charcot neuropathic osteoarthropathy Several studies have investigated associations between Charcot foot and the risk of lower extremity fractures. For example, Young and colleagues compared bone mineral density between individuals with diabetic peripheral neuropathy who also had Charcot foot and individuals with diabetic peripheral neuropathy who did not have Charcot foot [119]. They found that those who had Charcot foot had lower bone mineral density in the lower extremity compared to those who did not, and concluded that low bone mineral density may be a possible reason explaining an association between Charcot foot and increased risk of lower extremity fractures. In another study, Sinacore and colleagues investigated bone mineral density in the calcaneus among individuals with diabetic peripheral neuropathy and Charcot foot, individuals with diabetic peripheral neuropathy and no Charcot foot, and healthy controls [120]. They found the lowest calcaneal bone mineral density in individuals with diabetic peripheral neuropathy and Charcot foot.

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Areas of future biomechanical research

The prevalence of diabetic foot disease is increasing worldwide. Diabetic foot disease results in significant reductions in mobility, quality of life, and eventually life expectancy, and thus is a huge burden not only to individuals with diabetic foot disease but also to significant ones, care providers, and the health care system. In this chapter, we briefly review the biomechanical issues regarding the diabetic foot. Based on the current status of biomechanical research in the diabetic foot, the followings are some examples of future research. One primary issue is translating the findings from bench to bedside. The end product of the biomechanical understanding of the diabetic foot should be improvement in diabetic foot care. Although translating outcome measures might have been a challenge in the last few decades, with recent development of wearable technology and smart technology, monitoring some of the biomechanical measures during day-to-day activity is now possible [121,122]. In individuals with diabetic foot disease, translating biomechanical measures will be particularly important, given the relevance to ambulatory status. Another important issue is the development of novel yet practical perioperative biomechanical assessments. Although care providers for the diabetic foot have gained some insights from measuring plantar pressure, more accurate, efficient, and consistent measures have long been demanded [123]. This would significantly advance preoperative evaluation and postoperative management of diabetic foot. Another primary direction of future biomechanics research is to directly apply our biomechanical understanding to develop and improve current treatment paradigms in diabetic foot disease. A few studies demonstrated promising effects of foot and ankle exercise in diabetic foot disease. For example, Eraydin and Av¸sar investigated the effect of a 12-week foot and ankle exercise program on wound healing in people with diabetic foot disease and reported a significant reduction in mean ulcer areas [124]. Grewal and colleagues investigated a 4-week sensor-based interactive foot and ankle exercise on postural stability in people with diabetic peripheral neuropathy and demonstrated that this was significantly improved after 4 weeks [125]. Despite these promises, there still exists too little evidence to confirm the effect of foot and ankle exercise, and thus we recommend this as an important area for future biomechanics research. Additionally, we believe biomechanics research can contribute to develop emerging home-based therapies with respect to critical evaluation of these interventions. Recently, Kang and colleagues demonstrated plantar mechanical stimulation through a wearable foot compression device significantly improved plantar sensation, gait performance, and postural stability in people with diabetic peripheral neuropathy after 4 weeks [22]. Najafi and colleagues investigated the effect of home-based 6-week electrical stimulation on gait performance and postural stability in people with diabetic peripheral neuropathy and reported significant improvements in gait performance and postural stability after 6 weeks [126]. Researchers and clinicians have significantly advanced our understanding of the biomechanics of diabetic foot disease over the last few decades, and biomechanical research will undoubtedly continue to address the underlying mechanisms of diabetic foot disease and will contribute to the management of diabetic foot care through technological advances in the near future.

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Chapter 36

Rheumatic Foot Disease James Woodburn1, Ruth Barn2 and Gordon Hendry2 1

Griffith Centre of Biomedical and Rehabilitation Engineering (GCORE) and School of Health Sciences and Social Work, Griffith University, QLD,

Australia, 2School of Health and Life Sciences, Glasgow Caledonian University, Glasgow, United Kingdom

Abstract This chapter explores the biomechanics of the foot and ankle across important adult and childhood rheumatic diseases including rheumatoid arthritis, the seronegative spondylarthropathies, connective tissue disease, and juvenile idiopathic arthritis. The chapter draws on key literature and our own insights and experiences from clinical practice and research. Key biomechanical concepts are illustrated through selected patient case studies. In this clinical field, foot and ankle biomechanics is dominated by three-dimensional gait analysis research that describes pathological gait as a consequence of persistent primary disease mechanisms involving the joint, entheses, and tendons of the foot and ankle. We provide extensive descriptions of changes to spatio-temporal gait patterns as a product of impairment-driven compensation and destruction of foot and ankle joints. We explore how joint kinematics and kinetics, electromyographic muscle function, and plantar pressure distribution are disrupted across the life course of each rheumatic disease. We relate these observed changes to joint and soft tissue pathology, for example synovitis and enthesitis, that disrupt normal foot structure and function. We also attempt to explore the relationship between the patients’ lived experience in terms of impairments such as joint pain, stiffness and deformity, and ultimately progressive and irreversible disability, with biomechanical changes as detected and quantified via gait analysis. The chapter concludes with suggestions for future research that may lead to translational biomechanical breakthroughs to advance diagnoses and improve patient outcomes.

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Introduction

Rheumatic diseases predominantly affect joints, bones, tendons, ligaments, and muscles. They are commonly present in the foot and ankle; in rheumatoid arthritis (RA) for example, almost all patients will report foot-related problems during the disease course. Across these diseases, the interaction of biomechanics and primary disease mechanisms such as synovitis, where the synovial membrane that lines synovial joints becomes inflamed, determine the pattern and distribution of joint and soft-tissue damage. This in turn drives impairments such as pain, stiffness, and deformity that leads to functional changes and disability. Joint and soft-tissue inflammation play a major role across these conditions. Its initiation, localization, and persistence are influenced by biomechanical factors, particularly in the foot. For example, the anatomy and biomechanics of the Achilles tendon explain the localization of enthesopathic (pathology affecting insertion sites) lesions to bone, tendon, and bursa of the enthesis organ [1]. However, these conditions (outside of RA) are not well studied in the context of foot and ankle biomechanics. Significant attention has been paid to gait and how rheumatic diseases change foot structure and function during walking. In the last decade, technologies such as threedimensional gait analysis and multi-modal medical imaging have been successfully deployed to identify, characterize, and quantify changes in joint motion, kinetics, muscle function, plantar pressure distribution, and gait parameters that arise from foot and ankle pathology, as well as compensatory adaptation to pain, joint stiffness, and deformity. In this chapter, we explore the biomechanics of the foot and ankle in RA, the spondylarthropathies (a group of conditions sharing several common features including peripheral arthritis and enthesitis), juvenile idiopathic arthritis (JIA), gout, and connective tissue disorders. Attention will be paid to the interaction of primary disease mechanisms and biomechanics, subsequent damage to foot structures and function, and symptomology.

Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00001-9 © 2023 Elsevier Inc. All rights reserved.

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Rheumatoid arthritis

RA is a chronic inflammatory joint disease, which affects the joints and soft-tissues of the foot and ankle. It is characterized by joint inflammation, progressive joint destruction, and increase in functional disability. Cross-sectional studies in patients with established disease suggest that foot problems occur in around 80% 90% of patients, typically in the form of forefoot pain, stiffness, deformity, skin pressure lesions such as callosities, and problems related to footwear fit and comfort. Deformity can often be severe, involving the hind and mid-foot joints and presenting as pes planus. In the forefoot, the following signs and symptoms are common: general widening along with hallux valgus; hammer, claw or mallet toes; metatarsophalangeal joint subluxation and dislocation; plantar callosities and adventitious bursa; dorsal toe callosities; skin ulceration; and intermetatarsal neuroma. Extra-articular pathology including enthesitis (inflammation at the site where tendon attaches to bone), tendinopathy and bursitis (inflammation of fluid filled sacks in the joint) are also associated with biomechanical factors and appear alongside progressive and worsening deformities such as pes planus in the presence of tibialis posterior tenosynovitis (inflammation of the tendon sheath). Overall, the burden of foot disease in RA is substantial and impacts negatively on health-related quality of life [2].

36.2.1 Early rheumatoid arthritis Small joint inflammation in the hands and feet are the hallmark of early RA (, 2 years disease duration). Active joint disease leading to foot pain and walking disability has been reported in around 40% of patients at 2 years from diagnosis [3] and altered biomechanics are detectable. We investigated a cohort of early RA patients and found evidence of lesser toe deformities accompanied by a reduced weight-bearing contact and elevated peak pressures at the plantar metatarsophalangeal joints [4]. Plantar pressures are further elevated where the associated metatarsophalangeal joint showed erosive changes. If these pressures are not addressed, symptoms may persist or deteriorate and secondary pressure lesions such as callus, adventitious bursa, and ulceration are possible. Persistent synovitis in early RA leads to disabling forefoot pain and patients will adapt their gait to offload these painful sites. This is achieved by slowing walking speed and avoiding weight transfer to the forefoot. This compensation disrupts the rocker function of the foot and is characterized by: delay of transfer of center of pressure (CoP) to the forefoot; delay of heel rise and reduced peak vertical force; and reduced net ankle plantar flexor moment in terminal stance [4]. Patients may utilize a second compensation strategy through the adoption of a varus foot posture that leads to increased lateral loading, especially if the medial metatarsophalangeal joints are involved. Inflammatory synovitis and dysfunction of the peritalar joints and the tibialis posterior muscle-tendon unit are postulated mechanisms leading to instability of the hind and midfoot joints; clinically recognizable as pes planus. The natural history of this progressive deformity is not well understood but it can occur early and progress rapidly. Multi-segment foot models employed in 3D gait analyses have helped to dynamically characterize collapse of the medial longitudinal arch with accompanied increased hindfoot eversion and valgus pose during stance [5]. Collapse of the medial longitudinal arch creates a larger midfoot weightbearing surface and in early RA cases this has been shown to be up to 21% greater than healthy controls [4]. The third rocker function of the foot occurs in the latter part of the stance phase as the heel rises and this is controlled by concentric contraction of the ankle plantarflexors. Yet during manual muscle strength tests, the gastrocnemius-soleus muscle group is usually weak in people with RA. Consequently, ankle joint power is reduced during terminal stance and this can be detected in the early disease stages [4]. However, net muscular moments are observed to be normal, with reduced angular velocity at the ankle joint. Factors contributing to reduced angular velocity at the ankle joint such as walking speed and reduced range of motion (ROM) have been found in early RA patients but other factors including adaptations to impairments such as pain and deformity are considered important. Lastly, the vertical ground reaction force shows reduced loading rate, absent impact peak, and reduced first active peak during loading response suggesting that some early RA patients exercise caution in early stance, by slowing walking speed and decreasing stride length, to avoid painful symptoms in the hindfoot. [4].

36.2.2 Established rheumatoid arthritis In established RA, current models support an interaction of localized inflammatory and biomechanical factors that drive joint damage leading to degenerative and adapted gait and disability. This is mediated by distribution and severity of soft-tissue swelling, pain, joint stiffness, and deformity. With disease progression, patients with established RA walk slowly, have increased double support, and at the joint level, use smaller ranges of motion combined with reduced joint moments and power. Plantar pressures are abnormally distributed, especially in the forefoot associated with deformity (Figs. 36.1 and 36.2), and muscle function is disrupted and associated with tendinopathy. The importance of

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FIGURE 36.1 Peak plantar pressure distribution pattern (left) from a person with long-standing rheumatoid arthritis. Profile demonstrates absent or reduced hallux and lesser toe contact, high focal peak pressures at the first (960 kPa), second (445 kPa), and third (1055 kPa) metatarsal heads, and increased contact area in the medial longitudinal arch. The pattern is strongly related to the extensive forefoot impairments (right), including: hallux valgus; retraction and subluxation of the second and third metatarsophalangeal joints; and clawing of the fourth and fifth toes.

FIGURE 36.2 Peak plantar pressure distribution from person with long standing rheumatoid arthritis (left). In this 3D view the elevated peaks of pressure can be appreciated at the first, second, and third metatarsal heads and consistent with extensive forefoot impairments and plantar callosities over the first and third metatarsal heads (right).

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inflammatory synovitis over time reduces as structural damage accumulates. However, in the feet, pain of mechanical origin may persist as load bearing is carried through deformed feet and worsened further by difficult-to-fit footwear over prominent joints. In established disease, forefoot pain and deformity is associated with increased peak pressures in the forefoot, reduced contact area and a delay in the load transfer to the forefoot [6 8]. RA-associated changes to locomotor patterns are linked to adaptations to lower-limb segment coordination and reduced ROM, although patients can often efficiently adopt compensation mechanisms to overcome pathological changes to joint structures [9]. Adult-acquired pes planus deformity is common in patients with established RA. This condition is associated with significant localized pain and gait dysfunction, and it is difficult to treat, both conservatively and surgically. Pes planus is a clinical description of one or more geometric features, including loss of arch height, abduction of the forefoot, and valgus of the hindfoot. The planovalgus foot has a consistent pattern of multi-segment foot kinematics (Fig. 36.3). Radiologically, pes planus has been defined when standardized angles, such as the talocalcaneal angle (also known as the Kite angle), are above normal limits. The tarsal joints are relatively small and form complex structures, which are vulnerable to erosions, joint space narrowing and deformity associated with persistent synovitis. The exact mechanisms leading to pes planus in RA have yet to be elucidated, although several pathways linking inflammatory and biomechanical factors have been proposed. In vivo magnetic resonance imaging (MRI) based reconstruction of hindfoot kinematics in pathological hindfoot disease and in vitro cadaver-based models suggest that the normal motion-guiding and stability properties of ligaments and tendons surrounding the tarsal joints become compliant when inflamed joints are subjected to deforming loads on weight-bearing [10,11]. The role of tibialis posterior when tenosynovitic has been elucidated from 3D clinical gait analysis employing multi-segmented foot models and electromyography of tibialis posterior muscle [5]. An inflamed synovium, increased synovial fluid within the joint, joint capsule distension, and ligament damage across the joints combine to disrupt normal joint structure and function. In healthy subjects, under normal physiological loads, the largest tarsal joint rotations are found at the talonavicular joint, especially for coronal plane motion. This joint remains stable because the inferior calcaneonavicular ligament and the superomedial calcaneonavicular ligament are force-bearing and resist medial and plantar displacement of the talar head, assisted by the expansive insertion and blending of tibialis posterior to the tuberosity of the navicular. The talonavicular joint is vulnerable to synovitis in RA [2]. The superomedial portion of the calcaneonavicular ligament assumes the greater biomechanical role, but these structures are functionally interrelated, since calcaneonavicular ligament insufficiency is frequently encountered with tibialis posterior tendon dysfunction. Further, when tibialis posterior is dysfunctional, the midfoot loses its rigidity and stability during the latter part of stance [5]. The powerful gastrocsoleal complex then acts across the talonavicular joint as well as the forefoot during propulsion, and the resultant motion is thought to stretch the calcaneonavicular and medial plantar ligaments. Synovitis in the tibiotalar joint can disrupt the tibionavicular, anterior tibiotalar, and tibiocalcaneal portions of the medial deltoid ligament that stabilize tibial rotation and subtalar joint eversion [10]. The sinus tarsi is frequently painful and swollen in RA and synovitis leads to degeneration of the interosseous talocalcaneal and cervical ligaments, key stabilizers of the subtalar joint. Selectively attenuating these structures at or close to joints vulnerable to persistent synovitis in cadaveric specimens or kinematically analyzing feet with MRI-proven peritalar disease yields “rheumatoid arthritis like” pes planus type deformity in both approaches [11]. Tibialis posterior tendon disease occurs with sufficient frequency in RA to merit attention as a key factor in the development of pes planus [5]. The tibialis posterior muscle is contained within the deep posterior compartment with the tendon forming in the distal third of the leg. The tendon changes direction to enter the foot as it passes acutely behind the medial malleolus. In this region the tendon flattens and the tissue structure changes; exhibiting an increased presence of fibrocartilage and an avascular region. The location of the tendon relative to the axes of the subtalar and ankle joints facilitates inversion and plantarflexion respectively. Tibialis posterior is the most powerful invertor of the hindfoot as a result of the large inverter moment arm acting on the subtalar joint. The tendon has multiple insertions within the foot, dividing into three main components: (1) anterior, (2) middle, and (3) posterior. The anterior component is the largest and extends to the navicular tuberosity; it is reported to contain a fibrocartilaginous or bony sesamoid at this site which functions to provide a pressure absorbing or gliding mechanism. The middle and posterior components extend to the remaining tarsal bones, the middle three metatarsals and the flexor hallucis brevis muscle. The complex anatomy of the insertion sites functions to stabilize the medial longitudinal arch. As a classic enthesis its structure indicates a combination of tensile, compressive and shear forces acting on the insertion. Since the tendon sheath is part of the enthesis which is rich in fibrocartilage it is a primary site for tendinopathy in RA; and MRI and ultrasound studies have confirmed tenosynovitis, tendinosis, and attenuated structural changes in the form of thickening, degeneration, partial tear, and complete tear.

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FIGURE 36.3 (A) Multi-segmented foot joint 3D kinematics for (B) a person with rheumatoid arthritis (RA) and advanced foot impairments. In (C) and (D) data are presented as intersegment angle (deg) normalized for 0% 100% stance phase of gait. The colored line represents the motion pattern plotted for the person with RA against the black line with 1 / 2 1 standard deviation bars representing motion from otherwise healthy adults. In (C) data represent the hindfoot expressed in the coordinate system of the leg. Consistent with the clinical presentation, the hindfoot demonstrates reduced terminal stance extension, excessive eversion in the coronal plane and an internally rotated hindfoot position. In (D) the forefoot is expressed in the coordinate system of the hindfoot and shows increased terminal stance flexion, excessive inversion particularly in late stance and a forefoot external rotation position consistent with the planovalgus foot type.

A swollen and tender tibialis posterior tendon, particularly at the navicular insertion and retromalleolar regions accompanied by muscle weakness on manual testing, is indicative of tendon pathology. EMG has been employed, only rarely, to study tibialis posterior muscle function in RA: the muscle is only accessible via indwelling intramuscular electrodes. In healthly subjects, EMG activity for tibialis posterior shows a biphasic activity occurring during contact and either midstance or propulsive phases of gait with an amplitude of 20% 25% (when normalized by a maximum isometric reference contraction), although profiles are highly variable [12]. With RA, there is evidence to suggest

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FIGURE 36.4 (A) Adult acquired pes planovalgus in person with rheumatoid arthritis. (B) Everted pose of the hindfoot relative to the leg captured during midstance. (C) Hindfoot coronal plane motion normalized to 100% of stance. The hindfoot is everted at heel strike and continues to evert during the loading response ( . 15 degrees). Inversion is observed during terminal stance but insufficient to gain a neutral pose by toe-off. (D) The ankle joint coronal plane joint internal moment is predominantly opposing the everted position of the hindfoot during stance. (E) Consequently, power Doppler ultrasonography shows active tibialis posterior tenosynovitis and (F) EMG signal for tibialis posterior shows prolonged and elevated signal expressed relative to percentage of maximum voluntary contraction.

that tibialis posterior activity is amplified and prolonged during stance, especially in those patients with a pes planus phenotype [13]. EMG signal indicates loss of biphasic patterning with amplified activity in a monophasic pattern (through loading response, midstance, and load transfer). If EMG is combined with multi-segmented gait analysis these foot types also demonstrate predominantly invertor net muscular moments in the coronal plane that counteract the eversion motion of the hindfoot and plantar drift of the medial longitudinal arch (Fig. 36.4). Interestingly, in RA patients with confirmed tibialis posterior tenosynovitis and pes planus, custom foot orthoses designed to reduce the net invertor muscular moments can reduce tibialis posterior EMG amplitudes during stance, although the effect is highly variable [14]. Reducing walking speed is a well-recognized compensatory mechanism for pain in the foot, and this is independent of disease duration or activity. Joint destruction and deformity in the foot and ankle may further reduce walking speed irreversibly as joint damage burden accumulates. Walking speed therefore is strong quantitative determinant of both foot-related impairment and disability in RA [8]. Walking speed can be reliably measured in clinic over short distances and given the interdependency observed in bivariate analyses with joint angles, moments, and power it is a useful metric. For example, in foot intervention studies, lessening foot pain consistently leads to increased walking speed and this is accompanied by increased joint angles, joint moments, and power, collectively representing improved joint function. More generally, the Carroll et al. systematic review [15] also found ankle ROM was not significantly different in RA in comparison to healthy control subjects but ankle power was significantly decreased. Plantar pressure distribution demonstrated significantly increased forefoot peak pressures but this significance was lost for the hindfoot, midfoot, and individual metatarsal regions of interest. Large effect sizes were observed for all gait parameters where significant changes were found. Revolutionary advances in the medical management of RA have radically reduced joint destruction and irreversible disability. These advances include the introduction of biological therapeutics, such as tumor necrosis factor inhibitors, and a paradigm shift in terms of early aggressive management with targets of remission or low disease states [16]. Foot

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FIGURE 36.5 Mapping the individual spatiotemporal parameters for a person with inflammatory joint disease and foot impairments (combining pain, joint stiffness and deformity). Against gait laboratory age/gender matched normal values with colorcoded standard deviations (23 SD to 13 SD) this person demonstrates marked slowing of walking speed (21 SD), with reduced cadence (23 SD), prolonged cycle time (12 SD), normal stride length, and increased double-support time (11 SD).

and ankle involvement remains a problem, albeit there is the potential that fewer patients might present with fewer symptoms and progress slower than those in past decades [16]. Therefore, we can still draw on what we have learned from biomechanical analysis of foot and ankle disease in established RA, largely based on instrumented gait analysis. Presently, we still encounter patients with demonstrable changes to overall gait pattern, clinically characterized as slow and shuffling (Fig. 36.5). We reported spatiotemporal gait parameters in a cohort of 74 patients with established RA demonstrating marked reductions in walking speed, cadence, cycle time, and stride length with a prolonged doublesupport time [8]. A recent systematic review [15] meta-analyzed 31 papers reporting gait data for patients with RA. Our cohort-observed data was confirmed across 19 further studies with significant decreases in walking speed, cadence, and stride length with a lengthening of double limb support time. In established RA, synovitis and joint damage can also be observed in the midfoot, hindfoot, and ankle joints in between 25% and 60% of cases. To determine the biomechanical changes resulting from persistent disease, 3D motion capture has been employed since the mid 1990s with joint kinematics reported in around 20 studies. O’Connell et al. [17] were one of the first groups to demonstrate a relationship between pain and deformity of the forefoot, stance phase abnormalities in foot function, reduced walking speed, and moderate disability. We first employed multi-segmented foot models in 2004 to study the association between joint damage and small joint function [18]. Our model, based on the Oxford Multi-segment Foot Model [19], was developed to track the leg, hindfoot, and forefoot segments in three dimensions, and the hallux and navicular in one dimension. In a cohort of 11 RA patients with longstanding disease, abnormal foot motion was detected and, on a per-case basis, found to be closely associated with clinical impairments including pain, stiffness, and deformity [18]. For the hindfoot we found peak dorsiflexion, peak plantarflexion, and ROM were unchanged in RA in comparison to five healthy adults. The most notable difference occurred in the preswing period, where plantarflexion was both delayed and failed to pass neutral in the RA patients in comparison with normal, with an absolute difference in position of 11.6 degrees at toe-off. Hindfoot motion in the direction of inversion/ eversion was markedly different between the groups; the RA patients demonstrated moderately restricted hindfoot motion typically about an everted range throughout stance. Inversion was observed during terminal stance in the RA patients, but was insufficient to move the hindfoot through neutral into the inverted position adopted by the normal subjects. When expressed in the leg coordinate system, the hindfoot in RA patients functioned about an excessively externally rotated position, and this motion was highly variable in comparison with normal subjects. In the forefoot, the ROM was reduced for all three planes of rotation, suggesting that the forefoot tended to be stiffer in the RA patients. In the sagittal plane, dorsiflexion, plantarflexion, and total ROM were reduced in the RA patients. As in the hindfoot, forefoot plantarflexion during pre-swing was reduced (by 3.8 degrees) in the RA patients compared with normal subjects. In the coronal plane, peak inversion was reduced and peak eversion increased, with an overall reduction in the total ROM in the RA patients compared to normal subjects. In the transverse plane, peak adduction in the RA patients was reduced, peak abduction the same, and total ROM reduced in the RA patients in comparison with normal subjects. During full-foot contact the vertical position of the navicular was lower in the RA patients compared to normal; however, the observed difference in the peak height, observed at the start of terminal stance, was typically , 5 mm. Navicular height was more variable in the RA patients, and during terminal stance the maximum vertical displacement was lower in comparison with normal. Hallux flexion during the loading response and through midstance to pre-swing was reduced in comparison with normal values. Hallux extension was reduced in the RA patients during pre-swing, and the observed peak angle for extension and the ROM were much smaller in the patient cohort in comparison with normal. Over a series of further studies, these findings were extended to cohorts with severe foot involvement sub-grouped by localization of deformity (predominantly forefoot, hindfoot, or combined) [7]. RA patients with severe foot deformity were characterized by long-standing disease, moderate levels of disease activity, and moderate-high levels of

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foot-related impairment and disability. When grouped by predominant location of deformity, 3D gait analysis detected both common features and distinctly different subgroup patterns which mapped closely to the localized structural impairments, and featured elements of adapted and degraded gait patterns. Across the RA groups, patients tended toward a shuffling gait style characterized by slow walking speed and prolonged double-support. Sagittal plane kinematic data showed a delay in heel rise and a marked reduction in the magnitude of plantarflexion in terminal stance in all three RA groups in comparison to normative data, which suggests that the normal rocker function at the ankle and forefoot is lost [7]. Patients with predominantly forefoot disease avoid load transfer to the forefoot in response to pain and all three groups showed loss of third rocker function in terminal stance, emphasizing the universal involvement of the forefoot. Changes to hindfoot motion can be detected in most RA patients, notably at heel strike with excessive eversion through the entire stance phase. However, those with hindfoot valgus deformity demonstrate excessive eversion for longer periods of stance making it possible to differentiate between those with hindfoot and forefoot deformities. The pattern and range of hindfoot internal and external rotation in all the RA patients tends to be within normal limits, suggesting that the normal coupling motion between the hindfoot and leg is not always established. Lack of coupling may be a consequence of severe deformity accompanied with some disruption and increased compliance of ankle, hindfoot, and midfoot soft tissue structures and changes in joint geometry (as a consequence of joint pathology). Patients with hindfoot valgus deformity have a substantially higher number of abnormal kinematic features and appreciably greater levels of foot-related disability. This suggests that persistent synovitis and damage to the joints and soft tissue in the hindfoot has severe functional consequences, supporting earlier clinical studies [20]. These findings have been observed in other gait-based studies of RA patients awaiting foot surgery [21]. Further, association between structural pathology of joints and tendons and alteration of normal hindfoot, midfoot, and first metatarsophalangeal joint kinematics has been demonstrated elsewhere [22]. wFinally, we examined the impact of foot and ankle biomechanics in 74 RA patients with established disease using a minimum “core set” of biomechanical parameters. Here we attempted to identify independent biomechanical predictors of disease impact in terms of foot-related impairment and disability [8]. Our findings confirmed previous observations, notably RA patients walked slower and with prolonged double-support time. The initial foot contact angle in the sagittal plane indicated that the foot was placed to the ground more plantigrade and that heel rise in terminal stance was reduced. These patients demonstrated a significantly lower net plantarflexor moment and peak ankle power in comparison with able-bodied adults. The weightbearing heel alignment angle and peak hindfoot eversion in stance indicated a tendency toward valgus heel deformity accompanied by significant lowering of the medial arch. This was in conjunction with an increase in the midfoot contact area in RA patients. In the forefoot, RA patients demonstrated higher forefoot peak pressures whilst the toe contact area was markedly reduced. However, moderateto-large correlations were found among the biomechanical dataset such that walking speed emerged as strongest predictor for foot and ankle related impairment and disability. In the final regression model, foot disease burden (impairment and disability) could be independently predicted by foot pain, the number of swollen joints in the foot and walking speed indicating important drivers for disability formed by symptoms, primary disease mechanisms (inflammation), and biomechanics [8]. Since RA is recognizably characterized by foot deformity, foot pressure measurement techniques have been widely employed to detect and quantify associated plantar pressure distributions. Hallux and lesser toe deformities serve to reduce the forefoot contact areas and with migration of the plantar plate and fatty fibro padding, eroded and damaged metatarsophalangeal joints experience elevated localized pressures. Clinical research suggests that elevated forefoot plantar pressures are associated with underlying joint damage and pain during walking [4,7,23] or deformity [24,25], but not active disease (synovitis) [25]. A novel approach, measuring plantar pressure and shear stress simultaneously, detected higher shear-time integral and mediolateral shear stress in addition to peak pressure and pressuretime integral in patients with RA [26]. Sites of plantar pain also correlated to peak shear-time integrals. Increased plantar shear stresses was thought to result from localized bony erosions at metatarsophalangeal joints as well as pain-driven compensatory gait instability. It has been demonstrated that elevated forefoot peak pressures occur alongside reduced lesser toe contact area. In-shoe and platform pressure systems also permit analysis of the displacement of the CoP. This has been usefully employed to characterize forefoot off-loading as a predominant gait compensation strategy with loss of third rocker function in patients with RA. For example, Semple et al. [27] showed that in patients with RA the CoP transfer from the heel to forefoot is delayed with a shorter duration spent in, and increased velocity through, the forefoot. In summary, primary disease mechanisms in RA interact with biomechanics to drive the onset and progression of common impairments associated with pain, stiffness, and deformity and with that irreversible disability. Gait analysis has enabled a better understanding of the functional consequences of persistent disease in joints and soft tissues.

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Spondlyarthropathies

Foot and ankle biomechanics for the family of conditions known as spondylarthopathies, includes, but is not restricted to, psoriatic arthritis, ankylosing spondylitis (AS), Reiter’s syndrome, and reactive arthritis, and are largely under investigated. The exception perhaps is psoriatic arthritis which manifests with peripheral arthritis, dactylitis (inflammation of the toe), and enthesitis primarily involving the Achilles tendon and plantar fascia. Localized anatomy and biomechanics play an important role in disease pathogenesis [1,28]. The “enthesis organ” acts as the primary target of the inflammatory response in psoriatic arthritis [1]. Smooth transfer of forces between soft and hard tissues at the Achilles tendon is facilitated by the fibrocartilage-rich enthesis as well as the retrocalcaneal bursa and precalcaneal fat pad (also known as Kager’s fat pad) [29]. However, over a lifetime of repetitive mechanical loading, microtrauma is unavoidable. Consequently, in psoriatic arthritis, Achilles tendon “whole-organ” pathology occurs presenting as tendon thickening, focal erosions, and enthesophytes on the posterior calcaneal surfaces along with retrocalcaneal bursitis and diffuse osteitis. Mechanisms associated with Achilles tendon insertional angle change and calf muscle weakness affecting control of hindfoot motion have been hypothesized. The gastrocnemius and soleus are the main muscles forming the Achilles tendon and provide postural control and power during the propulsive phase of gait. The Achilles tendon crosses the ankle and subtalar joints and is subject to complex rotational movements and if calf muscle strength is weak then foot motion may be poorly controlled during loading. In particular, excessive hindfoot eversion changes the alignment of the calcaneus relative to the tibia and with it the insertional angle of the Achilles tendon. Increased dorsiflexion and eversion may change the tendon pulley function of the superior calcaneal tuberosity and with it the stress dissipating role of the retrocalcaneal bursa, precalcaneal fat pad, and the enthesis proper. This may trigger or accelerate microtrauma leading to the development of enthesopathic lesions. We investigated the influence of localized biomechanics at the Achilles tendon insertion in psoriatic arthritis and were unable to identify abnormal hindfoot motion among psoriatic arthritis patient groups with and without Achilles tendon enthesitis [30]. We did find that people with psoriatic arthritis and enthesitis subject their Achilles tendons to significantly lower stresses, as determined by estimated tendon force, yet possess the thickest tendons. This was best explained by the overall reduction in walking speed, a possible compensatory strategy to reduce heel pain. However, in a gait-based study we could only estimate total tendon force and not the distribution of load across the enthesis footplate [30]. Therefore, stress shielding as a putative mechanism cannot be ruled out. Achilles tendon loads have been approximated to three times body weight and over a lifetime microtrauma and damage are inevitable, irrespective of hindfoot motion, and this is evidenced by enthesophytes which can easily be found in otherwise healthy adults. In addition to the entheseal pathology within the foot, peripheral arthritis and dactylitis are also common in psoriatic arthritis. The severity and pattern of joint damage in psoriatic arthritis is much less than that observed in RA (in most cases). The metatarsophalangeal joints undergo similar pathological changes to those seen in RA specifically synovitis, joint subluxation, and erosive changes but psoriatic arthritis also demonstrates additional pathology such as new bone formation in the presence of erosions. We have reported high levels of forefoot involvement in psoriatic arthritis but this was not correlated with elevated plantar pressures [31]. We suggest that peripheral joint pain in psoriatic arthritis is related to both local and systemic factors. Dactylitis is a classic feature of psoriatic arthritis involving multiple structures with varying levels of severity and including flexor tenosynovitis and bone edema [32]. Despite being common in psoriatic arthritis, the functional consequences of dactylitis are not well understood. No differences in plantar pressure were reported when comparing those with a history of dactylitis to those without and compared to otherwise healthy individuals in a small cohort [32] but further work is required in the presence of active dactylitis. AS predominantly affects the axial skeleton resulting in bony ankylosis (stiffness) of the spinal column and increased kyphosis (outward curve) leading to reduced functional ability. The foot is frequently involved in AS and similar to psoriatic arthritis is prone to enthesitis of the Achilles tendon and plantar fascia. Despite disease modifying and biologic therapy, enthesitis and foot related impairment and disability remain prevalent [33]. Quantitative studies on the biomechanics of the foot and ankle in AS are lacking, a systematic review in 2015 of gait characteristics in inflammatory arthritis found only three papers concerning AS and only one of these studied the foot and ankle [15]. Gait adaptations in AS show a trend toward reduced walking speed and stride length and 3D motion analysis demonstrates reduced ankle joint plantarflexion with the loss of the third rocker function [34]. Much of the evidence surrounding the foot and ankle in AS pertains to the imaging features whereby pathology may be detected in the absence of clinical symptoms. In addition to the enthesitis, features such as bone erosions, joint space narrowing, subchondral sclerosis, and bone marrow edema have been reported [35] with the most common site of foot involvement in the hindfoot. Further work is required to more fully understand the biomechanical consequences of these pathological changes.

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36.4

PART | 6 Clinical Biomechanics of the Foot and Ankle

Juvenile idiopathic arthritis

JIA is an umbrella term for a group of heterogeneous disorders linked by persistent and unexplained synovitis for a period of at least six weeks in a child under 16 years. Importantly, JIA is not a juvenile version of adult RA and the patterns of active synovial disease can vary from monoarticular or oligoarticular (four joints affected or fewer), to symmetrical or asymmetrical polyarticular presentations. Persistent oligoarthritis is the most common disease subtype and typically involves an asymmetrical large joint presentation affecting the knee, ankle, and/or elbow. Arthritis in the lower limb is common in JIA at disease onset, with the knee affected most frequently (B70%) followed by the ankle (B35%) and subtalar (B12%) joints [36]. These prevalence rates are initially high and then stabilize with the initiation of medical therapies, with ankle joint synovitis prevalence decreasing to 6% 12% after 1-year post-diagnosis [36]. Midfoot and forefoot involvement is less common, affecting fewer than 5% of cases throughout the disease course [36]. Tenosynovitis and bursitis can occur at the hindfoot. Enthesitis and dactylitis affecting the lesser toes can also occur and can be debilitating, but these are subtype specific manifestations in enthesitis-related and psoriatic forms of JIA and have not been studied in detail. Early disease manifestations in JIA include pain, joint swelling, stiffness, and gait disturbances such as reduced walking velocity, increased double support time, and reduced step length driven largely by painful symptoms. JIA can follow an unpredictable course. Prolonged periods of persistent active synovitis predict “late stage” features of joint destruction characterized by radiographic progression including joint space narrowing, erosions, disturbed epiphyseal growth, and ankylosis [37]. These changes manifest as severe joint restrictions, fixed flexion deformities, and growth disturbances. The pathogenesis of foot and ankle deformities likely involve a sequence of raised intra-articular pressure from effusions, joint capsule distension, loss of ligament structural integrity, muscle weakness, loss of cartilage, and ultimately structural displacement of bone. Foot and ankle pain, deformity, and abnormal biomechanics can be present in the absence of inflammatory disease activity. Relatively little is known concerning the interaction between inflammation and biomechanical factors that contribute to degraded gait patterns and walking disability. Studies of lower limb biomechanics largely involve small heterogeneous samples of children with varying JIA disease subtypes at various disease stages. A significant proportion of early research in this field pre-dates modern day disease-modifying anti-rheumatic and biologic treatment paradigms. More recently, several studies have sought to narrow inclusion criteria to children with JIA who have foot and ankle involvement, which has served to improve our contemporary understanding of the impact of the disease. However, the nature of sampling frames adopted limit the ability to generalize the results to the entire JIA population. Ankle joint synovitis affects over 80% of patients with JIA within 5 years of diagnosis and is predictive of poor long-term outcome [38,39]. This leads to reduced joint ranges of motion, the development of varus or more commonly valgus deformities, as well as impaired mechanical function. Abnormal hindfoot sagittal plane motion has been well described in JIA [40,41] using 3D kinematic and kinetic gait analysis. Whilst the initial peak vertical ground reaction force typically does not deviate from normal, impairment of the third rocker function of the foot occurs during terminal stance. Impairment of the propulsion effect of the ankle has been characterized by a reduction of available ankle plantarflexion, reduced peak ankle dorsiflexion moment, reduced ankle power, and reduced second peak vertical ground reaction force [40,41]. This functional alteration is likely driven by compensation strategies such as reduction of walking speed and step length to avoid intra- and peri-articular ankle pain during extrinsic musculature contraction that is required to generate force in the propulsive phase of gait. Impairment may also be explained by reduced plantarflexor muscle strength which has been demonstrated in children with JIA [42]. The initial impairment from active joint disease may precipitate a spiral of deconditioning manifesting as loss of muscle strength and endurance as a result of pain and activity avoidance strategies, which can persist once the active inflammation has been medically resolved. There may also be significant reduction in available ankle plantarflexion due to progressive joint destruction which, with reduced moments, results in suboptimal conditions for adequate propulsion. Detailed evaluations of 3D multi-segment foot kinematics, kinetics, spatio-temporal parameters, and plantar pressure distribution have been undertaken in children with JIA who were in receipt of optimal medical management [43]. In the context of mild foot-related impairments, disability, and disease activity, only small changes in foot function consistent with degraded and adapted gait were detected. Observed changes involved trends toward increased midfoot dorsiflexion and reduced lateral forefoot abduction, within 3 5 degrees ranges. These findings were in contrast to past studies which have shown marked changes in foot function, characterized by reduced walking velocity and shortened step and stride length [40] altered load and pressure distribution characteristic of both high-arch and flat-foot postures [44], disturbance of the normal sagittal rocker functions [45], and muscle imbalance and reduced muscle strength [42,45]. This study confirmed clinical suspicions that foot posture and function is often preserved for children with JIA who are well managed medically [43]. On the contrary, children with particularly destructive disease subtypes and suboptimal medical histories can exhibit features that are indicative of planus and cavus foot types. Several potentially overlapping pathological

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mechanisms for pes planus in JIA have been postulated. Severe ankle joint disease resulting in ankle joint limitation may result in excessive dorsiflexion and hindfoot valgus rotations at the subtalar joint to enable tibial progression [41,44]. Anterior facet subtalar and talonavicular joint disease, resulting in calcaneonavicular ligament insufficiency may also result in a planus foot type as has been proposed in adult RA. Similarly, tibialis posterior tenosynovitis which has been observed in JIA using musculoskeletal ultrasound [46] may result in muscle dysfunction and inability of the medial longitudinal arch to withstand propulsive forces. These changes can be confirmed using plantar pressure analysis to detect an increase in midfoot contact area and kinematics to detect minimum navicular height and excessive peak hindfoot eversion. By comparison, the cavus foot type is a far less common presentation in JIA. Typical clinical signs include a varus hindfoot posture, a high medial longitudinal arch, stiff midtarsal joints, metatarsophalangeal hyperextension, and adducted forefoot relative to the midfoot. Reduced midfoot contact areas and higher lateral forefoot peak pressures and excessively high minimum navicular height can be confirmed using plantar pressure profiles [44] and multisegment foot kinematics. The mechanisms behind development of cavus foot posture are less clear, but are most likely driven by abnormal knee position and function characterized by external knee valgus moment, due to severe hip or knee disease [44]. In summary, foot and ankle structure and function can be significantly impaired in children with JIA who are not well managed medically, or who have disease that is resistant to medical management. Inflammatory disease activity affecting the hindfoot is common and is associated with pain and reduced functional capacity. Future gait analysis research in this population should seek to evaluate the impact of JIA on foot and ankle biomechanics in homogeneous subgroups.

36.5

Connective tissue disorders

Systemic lupus erythematosus is a complex multi-system disorder with a negative impact on health-related quality of life. The foot and ankle are frequently involved however quantitative data on the biomechanics of the foot and ankle are lacking. Objective assessment of the foot and ankle in systemic lupus erythematosus has only recently been undertaken for the first time and included spatio-temporal parameters, muscle forces, and plantar pressures [47]. These results indicated reduced walking velocity, reduced muscle strength (for hindfoot dorsiflexion, plantarflexion, inversion, and eversion), and reduced plantar pressure with higher pressure time integrals for individuals with this condition [47].

36.6

Gout

Gout is a common chronic disease of monosodium urate crystal deposition which has a predilection for the first metatarsophalangeal and ankle joints [48]. It is characterized by the onset of self-limiting but severely painful monoarthritic flares and periods of disease quiescence [49]. Acute arthritis affecting the first metatarsophalangeal joint is a hallmark feature, affecting around 73% of people with gout at least once over the course of their disease [50]. This typically presents as rapid onset of redness, heat, swelling, and extreme pain in the joint that is tender to touch. If treated sub optimally, affected joints can be subject to chronic synovitis, gouty tophi, tissue damage, and osteoarthritic changes [51]. Monosodium urate crystal deposition can also occur within tendons and ligaments, and the deposition patterns, which have an affinity for the Achilles tendon, are suggestive of biomechanical factors in gout pathogenesis [52]. Foot related pain, impairments and disability are common in both chronic and acute gout [53,54]. Established disease features such as the double contour sign, indicative of monosodium urate crystal deposition and tophus formation, that are detectable using musculoskeletal ultrasound are strong correlates of foot-related pain and disability [50]. People with gout tend to walk more slowly, with increased step and stance times [49,55]. These changes in gait likely occur initially as a pain avoidance compensation strategy, followed by deconditioning through impaired muscle activity and muscle atrophy in chronic disease. Studies employing barefoot and in-shoe plantar pressure analyses have found that people with gout exhibit reduced hallux peak pressures and increased midfoot pressure time integrals, which together are indicative of a compensatory first metatarsophalangeal joint off-loading strategy in terminal stance [54,55]. The decreased loading of the first metatarsophalangeal joint during the propulsive phase of gait results in reduced plantar flexor muscle activity which may result in loss of muscle strength [54]. Indeed, subsequent research utilizing a novel hand-held dynamometer protocol has demonstrated that people with gout exhibit reduced first metatarsophalangeal joint plantarflexion force relative to healthy matched controls [51] (Fig. 36.6). Reduced foot and ankle extrinsic muscle strength has been strongly associated with greater levels of foot pain and disability in people with gout [56]. In addition to the pain avoidance gait compensations theory, researchers have

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FIGURE 36.6 Double contour sign: Deposits of monosodium urate on the surface of the hyaline articular cartilage creates an echogenic (reflected ultrasound waves) line on the outer surface of the joint (white arrow).

proposed that the presence of tophi impacts negatively on normal tendon function by reducing tensile capacity [56]. The Achilles and peroneal tendons are most commonly affected by tophi infiltration and these tendons play a key role in deceleration and creating stability for propulsion [56]. Reduced ankle plantarflexion, inversion, and eversion muscle strength have been demonstrated in people with gout relative to healthy controls [56]. There is likely a sequence of events driven by painful symptoms, fear of flare and further pain avoidance, and reduced muscle function as a result of abnormal foot function, confounded by the presence of tophi deposition within the peri-articular ankle tendons. Abnormal foot function persists during asymptomatic periods when participants are not in flare. A key consideration of foot biomechanics in people with gout is the prevalence of comorbid conditions, particularly obesity. Obesity is a major independent risk factor for both gout and hyperuricemia [57]. A common finding across foot pain research in populations with obesity, gout, or hyperuricemia is the increase in midfoot pressure and/or contact area, which is indicative of a planus foot type [54,55]. A key distinction is that whilst midfoot loading is elevated across these populations, forefoot loading is increased in obese individuals and people with asymptomatic hyperuricemia, but reduced in people with gout [54,55]. As such, the combination of obesity and gout may magnify the interaction between mechanical and inflammatory factors that contribute to foot-related disability.

36.7

Future research

Systemic disease localization in the foot and ankle, particularly in conditions such as psoriatic arthritis, suggests breakthrough discoveries for the role of biomechanics in the initiation, and persistence of primary disease mechanisms such as synovitis, tenosynovitis, and enthesitis. This will require greater efforts toward soft-tissue biomechanics exploiting new capabilities in multi-modal medical imaging, experimental biomechanics, and mechanomics approaches. Heterogeneity in the onset, distribution and severity of foot and ankle involvement across the rheumatic diseases presents future research opportunities within the paradigm of precision medicine. In foot and ankle biomechanics, these future challenges can be met through deployment of novel techniques in computational modeling in person-specific musculoskeletal and finite element modeling as well as scaling up gait analysis, particularly in free-living conditions outside the laboratory. These discoveries are important with respect to translating new knowledge toward the development of novel, biomechanically based, interventions. Future research should address the development of primary and secondary biomechanical treatment targets as an important adjunct to pharmacological treatments to prevent irreversible joint damage and disability. For example, we do not fully understand pressure redistribution through custom foot orthoses for swollen and tender forefoot joints in people with RA. How much pressure relief is required to lessen or alleviate forefoot symptoms and what soft-tissue and joint biomechanical changes are this related to? Future research should also focus on the biomechanical mechanisms of footwear, orthotics, and surgical foot and ankle interventions thereby contributing to person-centered care. These approaches should be broad and generalizable over the rheumatic disease spectrum for all ages.

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[2] Helliwell P, Woodburn J, Redmond A, Turner D, Davys H. The foot and ankle in rheumatoid arthritis: a comprehensive guide. Edinburgh: Churchill Livingstone; 2007. [3] van der Leeden M, Steultjens MP, Ursum J, et al. Prevalence and course of forefoot impairments and walking disability in the first eight years of rheumatoid arthritis. Arthritis Rheum 2008;59:1596 602. [4] Turner DE, Helliwell PS, Emery P, Woodburn J. The impact of rheumatoid arthritis on foot function in the early stages of disease: a clinical case series. BMC Musculoskelet Disord 2006;7:102. [5] Barn R, Turner DE, Rafferty D, Sturrock RD, Woodburn J. Tibialis posterior tenosynovitis and associated pes plano valgus in rheumatoid arthritis: electromyography, multisegment foot kinematics, and ultrasound features. Arthritis Care Res 2013;65(4):495 502. [6] van der Leeden M, Steultjens M, Dekker JH, et al. 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[31] Hyslop E, McInnes IB, Woodburn J, Turner DE. Foot problems in psoriatic arthritis: high burden and low care provision. Ann Rheum Dis 2010;69(5):928. [32] Wilkins RA, Siddle HJ, Redmond AC, Helliwell PS. Plantar forefoot pressures in psoriatic arthritis-related dactylitis: an exploratory study. Clin Rheumatol 2016;35:2333 8. [33] Koca TT, Gogebakan H, Kocyigit BF, Nacitarhan V, Yildir CZ. Foot functions in ankylosing spondylitis. Clin Rheumatol. Dec 3. doi: 10.1007/ s10067-018-4386-6. [Epub ahead of print]. [34] Del Din S, Carraro E, Sawacha Z, et al. Impaired gait in ankylosing spondylitis. Med Biol Eng Comput 2011;49(7):801 9. [35] Erdem CZ, Sarikaya S, Erdem LO, Ozdolap S, Gundogdu S. MR imaging features of foot involvement in ankylosing spondylitis. Eur J Radiol Jan 2005;53(1):110 19. [36] Hendry GJ, Shoop-Worrall SJ, Riskowski JL, et al. 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Effects of juvenile idiopathic arthritis on kinematics and kinetics of the lower extremities call for consequences in physical activities recommendations. Int J Pediatr 2010;2010 pii: 835984. [41] Merker J, Hartmann M, Kreuzpointner F, et al. Pathophysiology of juvenile idiopathic arthritis induced pes planovalgus in static and walking condition: a functional view using 3D gait analysis. Pediatr Rheumatol Online J 2015;10(13):21. [42] Brostro¨m E, Nordlund MM, Cresswell AG. Plantar- and dorsiflexor strength in prepubertal girls with juvenile idiopathic arthritis. Arch Phys Med Rehabil 2004;85(8):1224 30. [43] Hendry GJ, Rafferty D, Barn R, Gardner-Medwin J, Turner DE, Woodburn J. Foot function is well preserved in children and adolescents with juvenile idiopathic arthritis who are optimally managed. Gait Posture 2013;38(1):30 6. [44] Fairburn PS, Panagamuwa B, Falkonakis A, et al. The use of multidisciplinary assessment and scientific measurement in advanced juvenile idiopathic arthritis can categorise gait deviations to guide treatment. Arch Dis Child 2002;87:160 5. [45] Truckenbrodt H, Hafner R, von Altenbockum C. Functional joint analysis of the foot in juvenile chronic arthritis. Clin Exp Rheumatol 1994;12: S10. [46] Hendry GJ, Gardner-Medwin J, Steultjens MPM, Woodburn J, Sturrock RD, Turner DE. Frequent discordance between clinical and musculoskeletal ultrasound examinations of foot disease in juvenile idiopathic arthritis. Arthritis Care Res (Hoboken) 2012;64:441 7. [47] Stewart S, Dalbeth N, Aiyer A, Rome K. Objectively-assessed foot and ankle characteristics in people with systemic lupus erythematosus: a comparison with age- and sex-matched controls. Arthritis Care Res (Hoboken) 2019;. Available from: https://doi.org/10.1002/acr.23832 [Epub ahead of print]. [48] Roddy E, Muller S, Rome K, et al. Foot problems in people with gout in primary care: baseline findings from a prospective cohort study. J Foot Ankle Res 2015;8:31. [49] Stewart S, Morpeth T, Dalbeth N, et al. Foot-related pain and disability and spatiotemporal parameters of gait during self-selected and fast walking speeds in people with gout: a two-arm cross sectional study. Gait Posture 2016;44:18 22. [50] Stewart S, Dalbeth N, Vandal AC, Allen B, Miranda R, Rome K. Are ultrasound features at the first metatarsophalangeal joint associated with clinically-assessed pain and function? A study of people with gout, asymptomatic hyperuricaemia and normouricaemia. J Foot Ankle Res 2017;10:22. [51] Stewart S, Dalbeth N, Vandal AC, Rome K. Characteristics of the first metatarsophalangeal joint in gout and asymptomatic hyperuricaemia: a cross-sectional observational study. J Foot Ankle Res 2015;8:41. [52] Dalbeth N, Kalluru R, Aati O, et al. Tendon involvement in the feet of patients with gout: a dual-energy CT study. Ann Rheum Dis 2013;72 (9):1545 8. [53] Rome K, Frecklington M, McNair P, Gow P, Dalbeth N. Foot pain, impairment, and disability in patients with acute gout flares: a prospective observational study. Arthritis Care Res (Hoboken) 2012;64(3):384 8. [54] Rome K, Survepalli D, Sanders A, et al. Functional and biomechanical characteristics of foot disease in chronic gout: a case-control study. Clin Biomech (Bristol, Avon) 2011;26(1):90 4. [55] Stewart S, Dalbeth N, Vandal AC, Rome K. Spatiotemporal gait parameters and plantar pressure distribution during barefoot walking in people with gout and asymptomatic hyperuricemia: comparison with healthy individuals with normal serum urate concentrations. J Foot Ankle Res 2016;30(9):15. [56] Stewart S, Mawston G, Davidtz L, et al. Foot and ankle muscle strength in people with gout: a two-arm cross-sectional study. Clin Biomech (Bristol, Avon) 2016;32:207 11. [57] Choi HK, Atkinson K, Karlson EW, Curhan G. Obesity, weight change, hypertension, diuretic use, and risk of gout in men: the health professionals follow-up study. Arch Intern Med 2005;165(7):742 8.

Chapter 37

The Aging Foot John B. Arnold1 and Hylton B. Menz2,3 1

Allied Health & Human Performance Unit, University of South Australia, Adelaide, SA, Australia, 2School of Allied Health, Human Services and

Sport, La Trobe University, Melbourne, VIC, Australia, 3La Trobe Sport and Exercise Medicine Research Centre, La Trobe University, Melbourne, VIC, Australia

Abstract Aging is associated with many changes to the musculoskeletal system, which over time becomes less well adapted to perform its normal functions. Alterations to the structure and biomechanics of bone, articular cartilage, ligament, tendon, fat, and skeletal muscles in the foot are a normal part of the aging process. They may also provide a basis upon which age-related diseases, such as osteoarthritis, can be initiated. The foot also contains a rich network of sensory receptors, nerves, lymphatic and circulatory vessels, all of which are also affected by aging. Combined with the demands of standing, walking, and other physical activities during a lifetime, foot deformity and functional impairment often develop. Understanding how aging affects the foot is essential for biomechanists and clinicians, allowing them to recognize the difference between normal and pathological states, how foot function changes over time, and how the response to treatments for foot pathology changes with advancing age. The aims of this chapter are the following:

1. to review the evidence regarding changes to the properties and functions of foot tissues with aging; 2. to examine the known age-related changes to foot anthropometry, posture, and morphology; 3. to explore how aging alters the function of the foot during walking, including joint kinematics, kinetics, and plantar pressures; 4. to review the evidence for how changes to foot posture and foot deformity with aging relate to mobility limitations in older adults; and 5. to identify and discuss areas for future research

37.1

Changing properties and functions of foot tissues

37.1.1 Bone Aging is associated with changes to the biomechanical properties of bone, underpinned by alterations to the amount, quality, and arrangement of its constituent minerals (mainly calcium hydroxyapatite), organic components (mainly type 1 collagen), and water [1]. By affecting both cortical and cancellous bone, aging results in reduced bone mass, decreased bone turnover, and decreased collagen content [2]. Reductions in bone mass occur from approximately the middle of the third decade of life [3,4] although the rate of change is dependent on gender, race, and anatomical site [5 7]. In trabecular bone, thinning and more perforations in the trabecular network, and a shift from a “plate-like” to “rod-like” structure occur (Fig. 37.1), with the most rapid changes occurring after 65 years of age [8,9]. Structurally, in the femur, the porosity of cortical bone increases and is a strong determinant of the reduction in strength, with decreases in ultimate stress (5%), strain (9%), and energy absorption (12%) per decade [10], when measured in both tension and compression [11]. For the metatarsals, reductions in bone bending strength are observed with aging [12,13], underpinned by reductions in areal bone mineral content and density. Reductions in bone mineral density (BMD; the average concentration of mineralized content in a defined section of bone) from adulthood in the axial and appendicular skeleton are also well recognized [6]. In the calcaneus, the rate of change in BMD accelerates with advancing age, from 1.2% per year from 65 69 years, to 2.6% per year after age 85 [14]. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00034-2 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 37.1 Tibial trabecular bone sample from 26-year-old male donor (left), compared to an 85-year-old female donor (right), showing changes from plate-like to rod-like structure and reductions in trabecular thickness. From Ding M, Hvid I, Quantification of age-related changes in the structure model type and trabecular thickness of human tibial cancellous bone. Bone 2000;26(3):291 95.

37.1.2 Cartilage Articular cartilage is an important regulator of the forces imparted on synovial joints during movement. Reduced cartilage thickness and volume [15], protein content [16], hydration [17], fibrillation of the superficial cartilage [18], increased matrix calcification [19], and formation of advanced glycation end products [20] are all features of aging joints. Over time, these changes to the cellular (chondrocytes) and extracellular matrix impair the ability of cartilage to withstand normal mechanical stress without damage [21,22]. The ankle joint has thinner cartilage compared to the knee and hip ( . 2 mm), with talar and tibial layers approximately 1 2 mm in thickness [23 26] and mean thicknesses for the subtalar, talonavicular, navicular-first cuneiform, calcaneocuboid, and first metatarsophalangeal joints all 1 mm or below [27,28]. Although it would appear logical that this would make smaller joints more susceptible to age-related diseases such as OA, primary ankle OA is much less common than knee OA and is usually post-traumatic in origin [29]. High joint congruency, as seen in the ankle, gives protection from mechanical overload by reducing peak stresses [26]. The type II collagen and protein content does not decrease with age in ankle cartilage [30] but this tissue does show an increased rate of collagen turnover and synthesis, higher stiffness, and better capacity for repair than cartilage in the knee [30,31]. Cartilage thickness is also related to the site and distribution of load across the articular surface in healthy adults, with the thickest cartilage in areas of high load [23]. Although the stiffness of cartilage in small midfoot joints such as the metatarsal-cuneiform joint is similar to other joints such as the ankle [32] it has been noted that the less columnar-like arrangement of chondrocytes suggests a reduced load-bearing function [32]. However, increased compressibility and permeability of cartilage in the midfoot and more vertical alignment of the articular surfaces suggest a propensity for a load-bearing function during dynamic activity, such as during the push-off phase of walking.

37.1.3 Muscle Aging is associated with a loss of skeletal muscle mass and cross-sectional area of up to 30% by age 75 compared to young adulthood [33]. Overall, the reduction in muscle strength of approximately 30% between 30 and 80 years of age is almost proportional to the loss of muscle mass [34,35]. However, widespread adaptations accompanying changes to muscle size are strongly related to reductions in muscle strength, beyond that imposed by reduced muscle tissue alone. Underlying gross morphological and strength changes with aging includes alterations in the total number of muscle fibers and the relative composition of fiber type (reduced fast glycotic, preservation of slow oxidative), a reduced size of the remaining fibers in the elderly, and reduced single-fiber tension [33,36]. Factors such as motor unit remodeling in the form of selective denervation of muscle fibers and a reduced number of motor units [37] and reductions in voluntary activation also contribute to the overall loss of muscle strength and power [38]. In the leg specifically, reduced cross-sectional area (16%) and volume (28%) of the gastrocnemius has been noted in elderly men, accompanied by reductions in voluntary activation (12%), specific fascicle force (30%), and muscle strength [39]. Age-related declines in plantarflexor and dorsiflexor maximal isometric muscle strength have been reported from the fifth and third decade of life, respectively [40]. Adults aged 70 79 had average dorsiflexion strength values that were 31% lower (females) and 25% lower (males) than their peak values achieved between 20 29 years of age. Losses

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FIGURE 37.2 Normative reference values for maximal isometric strength of the ankle dorsiflexors and plantarflexors, demonstrating age-related declines beginning from approximately the fifth decade of life. Adapted with permission from Wolters Kluwer Health, Inc.: McKay MJ, Baldwin JN, Ferreira P, Simic M, Vanicek N, Burns J and 1000 Norms Project Consortium. Normative reference values for strength and flexibility of 1,000 children and adults. Neurology 2017;88(1):36 43 [41].

of plantarflexion strength were slightly less, at 18% (females) and 22% lower (males) compared to the average peak values achieved at 40 49 years of age. In adults over 60, age was negatively correlated with plantarflexor (20.253) and dorsiflexor strength (20.333). Weaker or positive correlations in middle-aged adults (20 59 years) for dorsiflexors (20.135) and plantarflexors (0.165) is consistent with the inverted U-shaped relationship between age and strength across the lifespan (Fig. 37.2). Deficits in isometric muscle strength have also been observed for ankle invertors (29%), evertors (30%), and the intrinsic foot muscles that plantarflex the hallux (40%) and lesser toes (26.5%) when older adults are compared to young adults (mean 77 years vs 23 years) [42]. However, at present, our understanding of the mechanisms underlying reduced strength in the foot muscles is limited. In healthy adults, foot muscle size appears to reduce with age, with some data indicating the thickness and cross-sectional area of intrinsic foot muscles (i.e., abductor hallucis, flexor hallucis brevis, quadratus plantae, abductor digiti minimi, and flexor digitorum brevis) and toe flexors (flexor digitorum longus and flexor hallucis longus) is decreased by 19% 45% in older compared to younger adults [43]. Reductions in the strength of the hallux (39%) and lesser toe flexor muscles (35%) has been reported, with strength positively correlated with the size of the intrinsic toe flexor muscles. In related work, reduced cross-sectional area of abductor hallucis (211%), flexor hallucis brevis (223%), and quadratus plantae (214%) in older adults with toe deformities was observed compared to healthy controls [44].

37.1.4 Tendon Aging impacts not only the contractile component of the muscle-tendon unit but is also associated with changes in the structure and function of the tendon [45], making them more prone to tissue damage and injury [46,47]. In the foot, age-related anatomical variations in the Achilles tendon have been noted, with continuity of the fibers of the Achilles tendon and plantar fascia in neonates, which then become part of the periosteum in the calcaneal tuberosity by adulthood, creating two essentially separate structures [48]. Histological studies have noted a decrease in the diameter but a minor increase in the concentration of collagen fibrils in the Achilles tendon with age [49]. Analyses of tendon composition in vitro have revealed an increase in non-reducible collagen cross linking, reductions in extracellular water, polysaccharide content and stem cells, higher elastin content, and a reduction in collagen fiber crimp angle [50 52]. Most of these observations would equate to reduced tendon stiffness with aging, and limited evidence in vivo indicates that older tendons, namely the Achilles tendon, are more compliant and less stiff than younger tendons [53]. The magnitude of the difference in tendon stiffness between older (mean 68 years) and younger individuals (mean 24 years) was 36%, with stiffness also independently associated with single-leg balance duration and steadiness [53]. Other potential functional consequences of reduced tendon stiffness include reduced force generating capacity of the contractile component of the muscle-tendon unit due to a change in sarcomere working length [54] and slower force transmission from muscle to bone [55]. Amongst other factors such as reduced strength and balance, the reduced rate of force development may have implications for reacting quickly to unexpected slips or perturbations during movement. Despite the detrimental effects of aging on tendon, mechanical properties such as stiffness may be increased with resistance exercise [56,57], providing evidence that exercise may help attenuate age-related declines in tendon structure and function.

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37.1.5 Ligament Ligaments connect bone to bone and stabilize, guide and restrict joint motions. Primarily composed of type 1 collagen, elastin, and proteoglycans [58], ligaments primarily function to resist tensile load transmitted through the collagen fiber network [59,60]. Most of the research investigating age-related changes in ligaments has been performed on intraarticular types, such as the anterior cruciate ligament of the knee. Evidence from these studies suggests a reduction in the number of stem cells, disorientation, and mild degenerative changes of the collagen fiber network, increased fibrocartilage at ligament insertions, and fewer mechanoreceptors [61 64]. Biomechanical changes include decreased strength, stiffness, and energy absorption [65]. One cadaveric study found that aging was inversely correlated (r 5 20.39) with the tensile strength of the talofibular ligament in adults aged over 50 years [66]. Another small cadaveric study (n 5 9) noted that strain of the anterior talofibular ligament was greater for a given load in older compared to younger specimens, although no quantitative results were presented [67].

37.1.6 Skin Skin on the plantar surface of the foot plays an important role at the interface between the body and environment. Structural features unique to the plantar as compared to dorsal foot skin includes a thickened epidermis (B1.5 mm thick) [68], a more complex pattern of keratin expression [69], and an epidermal ridge [70]. The intricate ridges— which can be seen in footprints—amplify tactile sensitivity by directing forces from the external environment to more deeply located mechanoreceptors [70]. The high density of sudoriferous glands in glabrous skin may also function to optimize sweat levels and improve the frictional characteristics of skin on different surfaces [71]. Increased skin compliance with ridges may also offer improved ability to deform when compressive and shear stresses are applied, such as in walking [72]. Structural changes with aging in the superficial layer of the skin (epidermis), include slight thinning, decreased adherence of adjacent corneocytes, reduced water content in the stratum corneum, and flattening of the interface with the more deeply located dermis [73]. There are fewer eccrine (sweat) glands [74], indications of decreased dermal thickness [75], vascularity, and collagen and elastin fiber density [76 79]. Specifically in the foot, with aging there is an increase in the hardness and dryness of the plantar foot skin, reducing its elasticity and capacity for protection and healing [80,81]. The reduced functional capacity of the skin increases the propensity for conditions such as excessive dryness (xerosis), cracking, and the development of fissures, and the development of diffuse and localized thickening (corns and calluses) [82]. Overall, physiological and mechanical changes to the skin, both in the foot and elsewhere in the body, result in impaired barrier function and slower repair times after injury.

37.1.7 Neural Somatosensory input from the lower limb is recognized as an important source of sensory information involved in controlling limb movement and balance [83]. This includes proprioceptive information from muscle spindles and Golgi tendon organs around lower limb joints to relay information to the central nervous system regarding muscle length, tension, and joint angle [83,84]. Aging is associated with decreased muscle spindle diameter and number of intrafusal fibers [85,86], and with clinical manifestations including impaired joint position sense and kinaesthesia of the first metatarsophalangeal [87] and ankle joints [88,89]. The plantar foot is one of few sites in the body containing glabrous skin and contains four types of specialised cutaneous sensory receptors—known as mechanoreceptors—important for detection of vibration, pressure, and stretch (i.e., Meissner’s, Pacinian and Ruffini corpuscles, and Merkel’s disks). “Fast adaptive” receptors (Meissner’s & Pacinian corpuscles) comprise the majority of mechanoreceptors in the foot and respond to more rapid pressure changes, such as during dynamic activity [90]. Studies comparing older and younger adults have shown reduced density and distribution of mechanoreceptors in the plantar foot skin [70], resulting in decreased sensitivity to high frequency vibration and pressure [91 93] and poorer tactile acuity [94]. The formation of localized plantar skin calluses more often seen in older adults may contribute to decreased tactile sensitivity as thickened skin impedes the transmission of mechanical stimuli to mechanoreceptors located in the dermis [95,96]. As the foot forms the only contact between the body and ground during standing, cutaneous mechanoreceptors also play a role in modulating balance control [97]. Sensory and motor nerve conduction velocity also reduces with age [98,99] stemming partially from atrophy of myelinated sensory neurons [100,101] which has been linked to slower reaction times and walking speed observed with aging [102,103].

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37.1.8 Fat pad Underneath the plantar foot skin lies the specialised fibroadipose tissue (‘i.e., the fat pad’) which serves to anchor the skin to the foot bones, reduce stresses during weight-bearing, and protect the deeper neurovascular structures [104]. Accordingly, the fat pad is thickest at the regions of highest stress during weight-bearing such as the calcaneus, which is B20 mm thick when unloaded, deforming by 8 12 mm at peak stress during walking [105 107]. Aging does not seem to affect the thickness of the calcaneal fat pad; however, the viscoelastic properties are altered; it becomes more compressible and dissipates more energy when loaded [105]. The architecture of the heel fat pad varies across its depth; a superficial microchamber of dense connective elastic septal tissue interfaces with a deep macrochamber layer of densely packed adipocytes encased in thin elastin and collagen envelopes [104,108]. Owing to the differing composition and structure, the microchamber is up to 10 times stiffer than the macrochamber layer [109]. Estimated using shear wave ultrasound elastography, the increased stiffness with aging appears to mainly arise from the microchamber layer, with the deeper macrochamber layer unaffected [110]. However, whether apparent differences in stiffness with this technique translate to altered behavior of the heel pad with dynamic loading remains to be determined. Another site for increased fat pad thickness is the metatarsal heads, ranging from B9 mm thick under the first and third to 6 mm under the fifth metatarsal head in adults aged over 40 [106]. Similar to the calcaneal fat pad, thickness is maintained with aging [106]. Increased stiffness and similar strains at across different impact velocities suggest a poorer ability to attenuate stress and adjust to different weight-bearing loads [111,112]. This is pertinent as forefoot pain is associated with higher peak pressures underneath these regions during walking [113] suggesting a relationship between load attenuation in the soft tissues and pain. Toe deformities in older adults, such as claw toes, also displaces the fat pad from beneath the metatarsal head more distally to the sub-phalangeal region [114]. Elevation of the proximal phalanx increases the stiffness and reduces the amount of tissue in the submetatarsal fat pad [111] leading to elevated peak stresses under the forefoot during walking in healthy older adults [115] and those with type 2 diabetes [116].

37.2

Foot posture and morphology

The desire to quantify foot posture in clinical and research settings has driven the creation of a variety of measures over at least the past 50 years (for detailed review see Section 3 of this book). Alterations in foot structure and shape with aging have been assessed using a limited selection of these methods, mainly clinical measures of overall foot posture and linear or angular measurements using radiography and 3D surface scanning. Nonetheless, an array of age-related changes have been revealed in both overall foot posture, alignment, and shape of different regions and aspects of the foot.

37.2.1 Anthropometrics Measurement of anthropometric characteristics in different age groups suggests increases in ankle and instep circumference, forefoot width [117], and forefoot circumference ([118] 1105 adults; [117] 168 adults, 20 80 years, 24 in each decade) in older adults (aged over 65 years). Increases in circumferences and width may be explained by a higher prevalence of lower limb odema with aging [119] and age-related conditions such as type 2 diabetes and cardiovascular disease. Other reported changes include reduced medial and lateral ball length (distance from posterior heel to first and fifth metatarsal head, respectively) in older (65 1 ) compared to middle (45 64) and young adults (25 44 years [120]). However, some of these data suffer from a lack of normalization to account for differences in overall foot size between groups, leading to uncertainty in which particular differences are age-related [117,120]. Sampling bias arising from convenience-based methods may also not provide estimates representative of the wider population. Some of the agerelated changes in foot anthropometry may be gender specific, with differences between males and females noted using both traditional anthropometric and 3D foot scanning. Compared to men, women have a narrower heel width [119,121], smaller instep circumference [122,123], longer medial ball length [119,121], and greater hallux valgus angle [123,124].

37.2.2 Foot posture The assessment of overall foot posture in different age groups using simple clinical rating systems such as the Foot Posture Index [125] and aggregation of normative data [126] has identified a gradual increase in FPI scores (more flat foot posture) from middle age (Fig. 37.3).

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FIGURE 37.3 FPI-logit scores indicating overall foot posture vs age for 1007 adults without foot pathology. From Redmond AC, Crane YZ, Menz HB. Normative values for the foot posture index. J Foot Ankle Res 2008;1(1):6.

37.2.3 Arch height The height of the medial longitudinal arch gradually reduces with aging, with cross-sectional studies of different age groups showing lower arch heights in older compared to young adults [92] with changes from approximately the fourth decade of life [118]. Measurements from footprints also suggest age-related lowering of the medial arch, with an increase in the width and contact area of the midfoot region [92,127].

37.2.4 Joint range-of-motion Non-weight bearing ROM at the ankle joint complex has been assessed in different age groups using custom six degree-of-freedom measurement devices [128]. Declines in ROM each decade from 20 years of age are evident in most anatomical planes, for a total change between 20 and 80 years of 7 degrees (plantarflexion), 6 degrees (abduction), 7 degrees (adduction), and 3 degrees (inversion). Females tended to show greater reductions with age, with ankle complex ROM fairly stable with aging except for plantarflexion. Studies of advanced aging (70 years and above) show larger reductions in ankle plantarflexion in men compared to women [129], although the small sample size means these findings require further verification. Similar results for weight bearing ankle joint dorsiflexion have been observed, with an average reduction of 9 degrees between young (20 years) and older adults (80 years) [92]. For ROM of the first metatarsophalangeal joint (first MTPJ), larger age-related changes of 25 degrees were observed [92]. In the 1000 Norms Project, a large database of normative reference values, the average reduction for dorsiflexion from 20 29 years of age to 80 1 was 4 degrees (males) and 7 degrees (females) for dorsiflexion, and 5 degrees (males) and 3 degrees (females) for plantarflexion [40]. Findings also indicated weak negative correlations between age and ankle plantarflexion (2 2 0.119) and dorsiflexion ROM assessed in weight bearing (20.212) in adults aged over 60. Weaker or positive

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FIGURE 37.4 Normative reference values for maximum ankle dorsiflexion and plantarflexion demonstrating age-related declines across the lifespan. Adapted with permission from Wolters Kluwer Health, Inc.: McKay MJ, Baldwin JN, Ferreira P, Simic M, Vanicek N, Burns J and 1000 Norms Project Consortium. Normative reference values for strength and flexibility of 1,000 children and adults. Neurology 2017;88(1), 36 43.

correlations in middle-aged adults (20 59 years) for dorsiflexion (20.110) and plantarflexion (0.002) demonstrate that the age-related reductions occur most markedly after 60 years of age (Fig. 37.4).

37.3

Foot function (kinematics/kinetics/plantar pressures)

37.3.1 Kinetics 37.3.1.1 Ground reaction forces Aging is associated with reductions in the second peak of the vertical and peak anterior ground reaction forces in older adults at preferred walking speed [130 132]. The differences are amplified at faster walking speeds; Boyer, Andriacchi, and Beaupre [130] noted differences between young and older adults increased by 3% and 7% in the vertical and anterior GRF when compared at 1.6 m/s as opposed to 1.3 m/s. Similarly, Larish, Martin, and Mungiole [131] reported no significant differences in vertical or anterior GRF characteristics between age groups at 0.8 m/s but reductions in the second peak of the vertical GRF and anterior GRF of 5% and 19% at 1.3 m/s.

37.3.1.2 Joint moments and powers During normal walking, studies have demonstrated a reduced peak external ankle dorsiflexion moment [132] and ankle plantarflexor positive power during push-off in older adults [133,134]. Reduced ankle joint power generation in late stance may be explained by age-related changes in muscle-tendon properties, such as loss of skeletal muscle mass [135], altered pennation angle and fascicle lengths [136], reduced tendon stiffness [53], and smaller type II (fast-twitch) muscle fibers [137]. These changes have important consequences for the power and force generating capability of muscles for movement [138] and may be a contributing factor to the less propulsive gait pattern observed in older adults, illustrated by reductions in step length, stride length, and walking speed [139]. Evaluation of the external forces acting about the lower limb joints during walking has revealed a redistribution of joint moments in older adults (mean age 69 years) compared to young adults (mean age 21 years) [134]. This is most evident during terminal stance, with a lower peak external ankle dorsiflexion moment and power generation in older adults at both preferred and fast walking speeds [131,133,134,139]. Despite ankle muscle plantarflexor power being observed as a relatively strong predictor of step length in older adults (R2 5 0.52), other studies have revealed strategies to counteract the reduced contribution of the ankle joint to gait propulsion in terminal stance [133]. An increased external hip flexor moment and power generation just before toe-off suggests an increased reliance on hip flexor muscles generating a “pull-off” strategy for propulsion in older adults [134]. This also translates into increased positive work contributed by the hip to swing limb progression in pre-swing relative to decreased positive work at the ankle in terminal stance (Fig. 37.5).

37.3.1.3 Plantar pressures Loading of the plantar surface of the foot is altered with age, with cross-sectional studies of older adults finding lower peak pressure under the heel, midfoot, second to fifth metatarsal heads, and hallux during walking at preferred speed compared to young adults [92,141,142]. This is accompanied by lower maximum forces under the heel and hallux, and

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FIGURE 37.5 Change in positive work contributed by the hip joint to total positive work of the ankle and hip combined during the terminal stance and pre-swing phases of gait for young and older adults. Data from Winter et al. (1990) [140].

FIGURE 37.6 Mean peak plantar pressure during walking in the hindfoot, midfoot, and forefoot for different age groups. Data from 1000 healthy individuals [148].

lower force time integrals under the midfoot and hallux [92,142]. As the magnitude of plantar pressures and forces are speed dependent [142,143], reduced peak pressures and forces are partially explained by reduced walking speed and step length in older adults [92,144]. More recent findings from much larger normative datasets across the lifespan have revealed increased peak pressures under the forefoot with advancing age (Fig. 37.6) despite increases in contact time [145,146]. Important factors suggested to contribute to elevated peak forefoot pressures with aging include decreased ankle joint dorsiflexion range of motion (which is related to elevated forefoot loading) [147] and increased plantar soft tissue stiffness [105,110,111]. The center of pressure (CoP) characterizes the point of application of the resultant ground reaction force vector, revealing information about the location and transfer of force acting underneath the foot during movement. Changes in the CoP with aging indicate a more medially located trajectory with respect to the long axis of the foot during walking [149,150]. Specifically, the CoP is more medially located in older adults during loading response (0% 20% of stance), midstance (21% 50%), and terminal stance (51% 83%). Conversely, in pre-swing (84% 100% of stance) the CoP trajectory may be more laterally positioned in older adults [151]. A similar observation of increased lateral forefoot

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loading relative to the CoP trajectory has been observed in people with pathological restrictions in first MTPJ motion such as with osteoarthritis [152], suggesting more gradual age-related reductions in first MTPJ range-of-motion may explain subtle changes in CoP trajectory during walking. Both structural (i.e., foot posture, toe deformity) and functional (i.e., non-weight bearing joint range of motion) foot characteristics are related to the plantar loading characteristics of the foot in older people [144]. This implies an important role for age-related changes in foot structure and function to impact on functional ability with aging.

37.3.2 Kinematics 37.3.2.1 Ankle and foot Differences in ankle kinematics in older adults include a less dorsiflexed ankle at initial contact [132,133] less plantarflexed ankle at toe-off (3.6 6.9 degrees) [133,139], and reduced sagittal plane ROM [134,153]. Despite consistent evidence in the literature of changes to ankle joint function with age, and alterations to loading on the plantar foot, the kinematics of the foot itself has drawn little attention. One main reason is that historically the foot has previously been considered as a single rigid body for kinematic analysis with non-invasive methods [154]. Motion occurring within the foot is not considered with this approach. More recently the foot has been modeled as multiple segments [155] and subsequently one study has compared multi-segment foot kinematics between young and older adults during walking [156]. Independent of walking speed, older adults displayed a less plantarflexed calcaneus at toeoff (6.5 degrees), indicative of a less propulsive gait pattern. Older adults also had a smaller ROM of the midfoot (2.9 degrees) and smaller coronal plane ROM of the metatarsus (1.1 degrees) during stance compared to young adults. Taken along with findings from previous studies, this indicates reduced mobility within the foot, and at the ankle joint complex in older adults during walking.

37.4

Foot posture, foot disorders, and mobility limitations

As outlined earlier, aging is associated with a general tendency toward the development of a more planus foot posture and kinematic changes indicative of relatively everted foot function when walking. Although the functional implications of variations in foot posture and foot function in relation to lower limb overuse injury in younger and athletic populations remain uncertain [157,158], the literature relating to older people is more consistent and suggests that planus foot posture and everted foot function may partly explain the increased incidence of common structural foot disorders, foot symptoms, mobility limitations, and falls in older people.

37.4.1 Foot posture and foot deformity Several studies have reported associations between foot posture and common foot deformities in older people. An analysis of the US National Health Interview Survey of 74,721 adults conducted in 1990 found that self-reported “flat foot” was significantly associated with hammer toes and hallux valgus [159]. More recent studies using objective measures of foot posture and function derived from pressure platforms (the Framingham Foot Study and the Johnston County Osteoarthritis Project), similarly indicate that planus foot posture is associated with an increased likelihood of hammer toes and overlapping toes [160], while everted dynamic foot function is associated with hallux valgus and overlapping toes [160,161]. The mechanisms responsible for these associations are not well understood, but may relate to alterations in the function of muscles acting on the toes as a result of altered foot alignment and motion during gait [162,163].

37.4.2 Foot posture and foot symptoms Planus foot posture and everted foot function associated with aging may also increase the likelihood of developing foot symptoms. The Cheshire Foot Pain and Disability Survey of 3417 people in the UK reported that flat foot (determined by self-report) was associated with foot pain [164] while a cross-sectional postal survey of 2100 adults in Denmark found that self-reported planus foot deformity (based on line drawings) was significantly associated with foot pain in the past month [165]. Similar findings have been reported using objective measures of foot function in the Framingham Foot Study, where planus foot posture and everted foot function were associated with generalized pain in the foot as well as with heel and arch pain in older men [166]. As with the association between foot posture and structural foot disorders, the underlying mechanisms linking planus foot posture and everted foot function to foot pain remain unclear. However, cadaver studies have shown that simulating a flat foot by surgically releasing arch-supporting soft tissues

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results in increased plantar fascia strain, talonavicular joint motion, and dorsal compressive forces in the midfoot, factors that could potentially lead to tissue damage and development of foot symptoms [167,168].

37.4.3 Foot posture, mobility limitations, and falls Given the important role the foot plays in supporting the body when performing weightbearing activities and adapting to irregular terrain during when walking, planus foot posture also has the potential to influence mobility in older people and may contribute to an increased risk of falling. In a cross-sectional study of 305 community-dwelling older people Spink et al. [169] found that flat foot posture (measured using the Foot Posture Index) was a significant independent predictor of increased postural sway when standing on a compliant surface. Similarly, a cross-sectional analysis of the Framingham Foot Study [170] found that compared to those with normal foot posture and function, participants with planus foot posture were more likely to report difficulty in remaining balanced, and those with everted foot function were more likely to report difficulty in walking across a small room. An analysis of the same population by Awale et al. [171] found that older people with planus foot posture had 78% higher odds of reporting two or more falls in the previous 12 months. Further evidence to support the association between planus foot posture, postural instability, and falls can be derived from laboratory-based studies assessing balance performance when wearing arch-supporting foot orthoses, and clinical trials reporting reductions in falls when older people are provided with orthoses as part of a multifaceted intervention. Gross et al. [172] reported significant improvements in single leg stance, tandem stance, tandem gait, and alternating step tests when older people wore arch-supporting orthoses with medial wedging to limit foot eversion. Similarly, Chen et al. [173] reported significant reductions in bipedal postural sway when older people wore foam arch supports with a heel cup. Finally, one randomised trial reported improvements in the Berg Balance Scale and Timed Up and Go Test in older women with osteoporosis who were prescribed custom foot orthoses made of ethylene-vinyl acetate with a medial arch support [174], and three randomized controlled trials have reported significant reductions in the incidence of falls in older people allocated to a multifaceted intervention consisting of prefabricated foot orthoses, footwear advice, and a foot and ankle exercise program [175 177].

37.5

Areas for future research

Our understanding of the biomechanics of the aging foot ultimately has implications for the design of effective treatments for foot pain and deformity in older adults. For financial and pragmatic reasons, our understanding of age-related changes to foot biomechanics is largely based on small cross-sectional studies comparing young and older adults. Prospective, longitudinal studies would improve our understanding of how aging affects the structure and function of the foot, without the confounding effect of inter-individual differences. What is considered “normal” aging as opposed to the effect of comorbidities is also an ongoing issue in gerontology research [178]. This is particularly important given that the same mechanisms that drive aging also drive many age-related chronic diseases, making it difficult to differentiate, from a mechanistic standpoint, what is truly a biological effect of aging alone [179]. While some studies have investigated the relationship between selected biomechanical variables (i.e., peak pressure), structural measures (i.e., metatarsal lengths), and the current or past forefoot pain [113], the link between changes in soft tissues, such as forefoot and heel pat pads, and symptoms has not been fully established. Forefoot pain is common in older people, affecting approximately 25% of adults aged over 65 years of age [180], highlighting the need for a better understanding of factors related to the development of forefoot pain in this age group. Subsequently, this may aid in the development and selection of materials, footwear, and foot orthoses to more effectively alleviate forefoot pain in clinical settings. Whether foot pain and deformity can be prevented by addressing specific age-related biomechanical alterations to foot structure and function is unknown and would also have implications for improved care of the health of older adults. The body’s response to exercise, foot orthoses, and footwear is an emerging area of research. To date, the effect of interventions on foot biomechanics has largely been focused on the periphery. There is scope to direct attention to the central nervous system to understand how interventions act to effect pain and movement control of the foot. Given that aging affects a plethora of physiological processes and systems, it would seem prudent to investigate the response of all parts of the musculoskeletal system to interventions. Advances in techniques such as weight-bearing computed tomography [181], diagnostic ultrasound [182,183], and electromyography [184] provide insight into the state and biomechanical response of neural and musculoskeletal tissues. Further studies applying these techniques at the nexus between aging, pathology, and interventions will continue to advance our understanding of the aging foot.

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[140] Winter D, Patla A, Frank J, Walt S. Biomechanical Walking Pattern Changes in the Fit and Healthy Elderly. Physical Therapy 1990;70 (6):340 7. [141] Hessert M, Vyas M, Leach J, Hu K, Lipsitz L, Novak V. Foot pressure distribution during walking in young and old adults. BMC Geriatr 2005;5(1):8. [142] Kernozek T, LaMott E. Comparisons of plantar pressures between the elderly and young adults. Gait Posture 1995;3(3):143 8. [143] Warren GL, Maher RM, Higbie EJ. Temporal patterns of plantar pressures and lower-leg muscle activity during walking: effect of speed. Gait Posture 2004;19(1):91 100. [144] Menz HB, Morris ME. Clinical determinants of plantar forces and pressures during walking in older people. Gait Posture 2006;24(2):229 36. [145] Bosch K, Nagel A, Weigend L, Rosenbaum D. From “first” to “last” steps in life pressure patterns of three generations. Clin Biomech 2009;24(8):676 81. [146] McKay MJ, Baldwin JN, Ferreira P, Simic M, Vanicek N, Wojciechowski E, et al. Spatiotemporal and plantar pressure patterns of 1000 healthy individuals aged 3 101 years. Gait Posture 2017;58:78 87. [147] Morag E, Cavanagh P. Structural and functional predictors of regional peak pressures under the foot during walking. J Biomech 1999;32 (4):359 70. [148] McKay MJ, Baldwin JN, Ferreira P, Simic M, Vanicek N, Wojciechowski E, et al.1000 Norms Project Consortium Spatiotemporal and plantar pressure patterns of 1000 healthy individuals aged 3 101 years. Gait Posture 2017;58(10):78 87. [149] Chiu M-C, Wu H-C, Chang L-Y, Wu M-H. Center of pressure progression characteristics under the plantar region for elderly adults. Gait Posture 2013;37(3):408 12. [150] Hagedorn TJ, Dufour AB, Golightly YM, Riskowski JL, Hillstrom HJ, Casey VA, et al. Factors affecting center of pressure in older adults: the Framingham Foot Study. J Foot Ankle Res 2013;6(1):18. [151] Sole G, Pataky T, Sole CC, Hale L, Milosavljevic S. Age-related plantar centre of pressure trajectory changes during barefoot walking. Gait Posture 2017;57:188 92. [152] Menz HB, Auhl M, Tan JM, Buldt AK, Munteanu SE. Centre of pressure characteristics during walking in individuals with and without first metatarsophalangeal joint osteoarthritis. Gait Posture 2018;63:91 6. [153] Hageman P, Blanke D. Comparison of gait of young women and elderly women. Phys Ther 1986;66(9):1382 7. [154] Cappozzo A, Catani F, Della Croce U, Leardini A. Position and orientation in space of bones during movement: anatomical frame definition and determination. Clin Biomech 1995;10(4):171 8. [155] Deschamps K, Staes F, Roosen P, Nobels F, Desloovere K, Bruyninckx H, et al. Body of evidence supporting the clinical use of 3D multisegment foot models: a systematic review. Gait Posture 2011;33(3):338 49. [156] Arnold JB, Mackintosh S, Jones S, Thewlis D. Differences in foot kinematics between young and older adults during walking. Gait Posture 2014;39(2):689 94. [157] Dowling GJ, Murley GS, Munteanu SE, Smith MM, Neal BS, Griffiths IB, et al. Dynamic foot function as a risk factor for lower limb overuse injury: a systematic review. J Foot Ankle Res 2014;7(1):53. [158] Neal BS, Griffiths IB, Dowling GJ, Murley GS, Munteanu SE, Franettovich Smith MM, et al. Foot posture as a risk factor for lower limb overuse injury: a systematic review and meta-analysis. J Foot Ankle Res 2014;7(1):55. [159] Shibuya N, Jupiter DC, Ciliberti LJ, VanBuren V, La Fontaine J. Characteristics of adult flatfoot in the United States. J Foot Ankle Surg 2010;49(4):363 8. [160] Hagedorn TJ, Dufour AB, Riskowski JL, Hillstrom HJ, Menz HB, Casey VA, et al. Foot disorders, foot posture, and foot function: the Framingham foot study. PLoS One 2013;8(9):e74364. [161] Golightly YM, Hannan MT, Dufour AB, Hillstrom HJ, Jordan JM. Foot disorders associated with overpronated and oversupinated foot function: the Johnston County Osteoarthritis Project. Foot Ankle Int 2014;35(Jul 18):1159 65. [162] Glasoe WM. Treatment of progressive first metatarsophalangeal hallux valgus deformity: a biomechanically based muscle-strengthening approach. J Orthop Sports Phys Ther 2016;46(7):596 605. [163] Green DR, Brekke M. Anatomy, biomechanics, and pathomechanics of lesser digital deformities. Clin Podiatr Med Surg 1996;13(2):179 200. [164] Garrow AP, Silman AJ, Macfarlane GJ. The Cheshire Foot Pain and Disability Survey: a population survey assessing prevalence and associations. Pain 2004;110:378 84. [165] Molgaard C, Lundbye-Christensen S, Simonsen O. High prevalence of foot problems in the Danish population: a survey of causes and associations. Foot (Edinb) 2010;20(1):7 11. [166] Menz HB, Dufour AB, Riskowski JL, Hillstrom HJ, Hannan MT. Association of planus foot posture and pronated foot function with foot pain: the Framingham foot study’. Arthritis Care Res (Hoboken) 2013;65(12):1991 9. [167] Kitaoka HB, Luo ZP, An KN. Mechanical behavior of the foot and ankle after plantar fascia release in the unstable foot. Foot Ankle Int 1997;18(1):8 15. [168] Kogler GF, Solomonidis SE, Paul JP. In vitro method for quantifying the effectiveness of the longitudinal arch support mechanism of a foot orthosis. Clin Biomech (Bristol, Avon) 1995;10(5):245 52. [169] Spink M, Fotoohabadi MR, Wee E, Hill KD, Lord SR, Menz HB. Foot and ankle strength, range of motion, posture, and deformity are associated with balance and functional ability in older adults. Arch Phys Med Rehab 2011;92(1):68 75. [170] Menz HB, Dufour AB, Katz P, Hannan MT. Foot pain and pronated foot type are associated with self-reported mobility limitations in older adults: the Framingham Foot Study. Gerontology 2016;62(3):289 95.

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[171] Awale A, Hagedorn TJ, Dufour AB, Menz HB, Casey VA, Hannan MT. Foot function, foot pain, and falls in older adults: the Framingham Foot Study. Gerontology 2017;63(4):318 24. [172] Gross MT, Mercer VS, Lin F-C. Effects of foot orthoses on balance in older adults. J Orthop Sports Phys Ther 2012;42(7):649 57. [173] Chen TH, Chou LW, Tsai MW, Lo MJ, Kao MJ. Effectiveness of a heel cup with an arch support insole on the standing balance of the elderly. Clin Interv Aging 2014;9:351 6. [174] de Morais Barbosa C, Barros Bertolo M, Marques Neto JF, Bellini Coimbra I, Davitt M, de Paiva Magalhaes E. The effect of foot orthoses on balance, foot pain and disability in elderly women with osteoporosis: a randomized clinical trial. Rheumatology (Oxford) 2013;52(3):515 22. [175] Cockayne S, Adamson J, Clarke A, Corbacho B, Fairhurst C, Green L, et al. Cohort randomised controlled trial of a multifaceted podiatry intervention for the prevention of falls in older people (The REFORM Trial). PLoS One 2017;12(1) (no pagination), e0168712, January. [176] Spink MJ, Menz HB, Fotoohabadi MR, Wee E, Landorf KB, Hill KD, et al. Effectiveness of a multifaceted podiatry intervention to prevent falls in community dwelling older people with disabling foot pain: randomised controlled trial. BMJ 2011;342:d3411. [177] Wylie G, Menz HB, McFarlane S, Ogston S, Sullivan F, Williams B, et al. Podiatry intervention vs usual care to prevent falls in care homes: pilot randomised controlled trial (the PIRFECT study). BMC Geriatrics 2017;17(1):143. [178] Lash TL, Mor V, Wieland D, Ferrucci L, Satariano W, Silliman RA. Methodology, design, and analytic techniques to address measurement of comorbid disease. J Gerontol Ser A: Biol Sci Med Sci 2007;62(3):281 5. [179] Fabbri E, Zoli M, Gonzalez-Freire M, Salive ME, Studenski SA, Ferrucci L. Aging and multimorbidity: new tasks, priorities, and frontiers for integrated gerontological and clinical research. J Am Med Dir Assoc 2015;16(8):640 7. [180] Hill C, Gill T, Menz H, Taylor A. Prevalence and correlates of foot pain in a population-based study: the North West Adelaide health study. J Foot Ankle Res 2008;1(1):2. [181] Barg A, Bailey T, Richter M, de Cesar Netto C, Lintz F, Burssens A, et al. Weightbearing computed tomography of the foot and ankle: emerging technology topical review. Foot Ankle Int 2018;39(3):376 86. [182] Kasehagen B, Ellis R, Pope R, Russell N, Hing W. Assessing the reliability of ultrasound imaging to examine peripheral nerve excursion: a systematic literature review. Ultrasound Med Biol 2018;44(1):1 13. [183] Shiotani H, Yamashita R, Mizokuchi T, Naito M, Kawakami Y. Site-and sex-differences in morphological and mechanical properties of the plantar fascia: a supersonic shear imaging study. J Biomech 2019;85:198 203. [184] Ferrari E, Cooper G, Reeves N, Hodson-Tole E. Surface electromyography can quantify temporal and spatial patterns of activation of intrinsic human foot muscles. J Electromyogr Kinesiol 2018;39:149 55.

Chapter 38

Biomechanics of Athletic Footwear Gillian Weir and Joseph Hamill Biomechanics Laboratory, Department of Kinesiology, University of Massachusetts Amherst, Amherst, MA, United States

Abstract The development of sports-specific shoes began in the 1970s when the explosion in popularity of running for physical activity began. The biomechanical evaluation of footwear is focused on two guiding principles for footwear design: (1) decreasing the risk of running-related injuries; (2) improving performance, with the former principle garnering the most attention. Impact loading and hindfoot stability were the most studied areas of running injuries. Despite the extensive research in footwear design and subsequent footwear innovations, running injury incidence remains unchanged. This chapter presents an overview of parameters used to guide footwear design, types of athletic and casual footwear, discuss shod versus barefoot running, describe footwear-related injuries and, finally, present some new paradigms for future footwear evaluation to reduce running-related injuries.

38.1

Introduction

Footwear has always held a special significance for humans as symbols of wealth and status or as simple protection from severe weather or intense terrain. According to anthropologists, the first modern humans (i.e., homo sapiens) walked out of Africa between 100,000 and 200,000 years ago moving north into the colder, more inclement climates of Europe and Asia. Then, a need arose for foot coverings to protect the feet from the environment. Researchers believe that the first shoes were constructed during the Ice Age although foot coverings may have been in use as early as 500,000 years ago in the northern climates by other early hominids such as Neanderthals and homo erectus [1]. However, the idea of protective footwear came much later and shoes were not in widespread use until approximately 26,000 40,000 years ago [2]. Few examples of early footwear have been found although it is surmised that the ability of early humans to make footwear was available. The earliest footwear that have been found dates back to the last Ice Age or about 10,000 years ago. An example of such footwear was discovered in Fort Rock, Oregon [3]. Researchers cannot state for certain why humans stopped going barefoot and developed footwear but the use of shoes is seen in the skeletons of that time period. At the beginning of the Tudor period in England (1485), most basic footwear construction techniques were developed and standardized. By the 17th century, high heels became fashionable for men and flat shoes with criss-crossed ribbons in imitation of classical sandals were a mainstay for women. By the 1800s, running races were quite common although specialized footwear was not considered necessary. However, in 1839, a significant development occurred; Goodyear developed a process using rubber in different contexts and, by 1850, rubber shoes were used for running. The first spike shoes were developed, not for running, but for cricket and croquet. The first spike running shoes were created in Northampton, England around 1850 [1]. By the early 1900, athletic footwear was mainly produced by rubber companies. US Rubber introduced Keds in 1916 while Converse introduced the first mass marketed basketball sneaker in 1917 with their introduction of the Chuck Taylor All Star. Other companies, including B.F. Goodrich and Spalding Co., were producing tennis shoes and other smaller companies were manufacturing early cleated shoes. After World War I, America turned to sports and physical activity and the market for athletic footwear grew steadily. In the 1920 and 1930, manufacturers added traction to the soles of their athletic footwear and began marketing them for different sports. A major innovation of this time was the production of distinct models for boys and girls, specifically, narrower lasts for girls. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00006-8 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 38.1 Evolution of athletic footwear.

FIGURE 38.2 Running shoe anatomy.

Footwear, in particular athletic footwear, continued development throughout the first half of the 20th century (Fig. 38.1). In the 1970s, jogging became a popular sport, and a need for specialized shoes was created. The specialization of athletic footwear, that is, specific footwear for specific activities, driven by the science of footwear, became the standard. Thus, the design of footwear was significantly influenced by researchers, particularly in the area of biomechanics.

38.2

Anatomy of a running shoe

There are basic structural elements to every shoe whether it be running, walking, hiking, work, or casual shoe (Fig. 38.2). However, not all of the elements are present in every type of shoe. These structures include: (1) outsole; (2) midsole; (3) insole board; (4) sock liner; (5) tongue; (6) heel counter; (7) upper; (8) collar; and (9) lacing system. These elements generally have a specific function and the elements are designed to accommodate this function. For example, the outsole (the element of the shoe that contacts the surface of the ground) can be designed with “nubs” to enhance traction. If shock attenuation is necessary, the midsole of the shoe, interacting with the outsole and the sock liner, can reduce the impact shock of the foot/ground collision.

38.3

Biomechanics of athletic footwear design

The evaluation and eventual design and construction of footwear have been primarily focused on two guiding principles: (1) decreasing the risk of footwear related injuries; and, to a lesser extent, (2) improving performance. The majority of research in footwear, from a biomechanical perspective, has been in decreasing the risk of injury. In terms of

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FIGURE 38.3 Features of the vertical ground reaction force (vGRF) used in the evaluation of footwear.

FIGURE 38.4 (A) Midsole durometer tested on an impact machine. (B) Midsole durometer tested on participants running across a force platform.

reducing the probability of footwear as a risk factor for locomotion-related injuries via means of footwear design, two areas have received much of the focus: (1) cushioning (i.e., reducing the load on the body from the foot/ground collision); and (2) maintaining hindfoot stability (i.e., preventing hindfoot eversion, medial longitudinal arch collapse, and forefoot external rotation).

38.3.1 Cushioning The measurement of external loads applied to the body and the attenuation of these loads via footwear has primarily focused on ground reaction force (GRF) data. More specifically, running research has primarily investigated the first peak force (active peak) and the loading rate of the vertical component of the GRF (Fig. 38.3). It was thought that these parameters were critical factors in reducing the risk of impact-type injuries. The relationship between impact forces and injury was initially derived from animal studies in which lower extremity joints of the animals were subjected to numerous repeated impacts. In a study by Radin and colleagues [4], joints were significantly degraded suggesting that the repeated impulsive impacts were deleterious. In the early period of the involvement of footwear biomechanists in shoe design, it was surmised that the way to reduce impact forces was to make the midsole of the shoe much softer. However, it was determined that softer materials did not necessarily result in better cushioning based strictly on the impact loading as measured from the vertical GRF. The soft midsole of the shoe could, in fact, collapse to the point whereby the vertical GRF was similar to a firmer midsole. This may be due to individuals adapting to the magnitude of the force when running [5]. Midsoles with varying hardness were tested on a mechanical impact tester (Fig. 38.4A). These same insoles were tested with human participants running in them

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illustrating that the impact peak exhibits very little difference in three types of shoes with varying midsole hardness (Fig. 38.4B). The inference is that that softer materials do not necessarily result in better cushioning and subsequently there is no reduction in the external forces applied to the runner. Newer evidence now suggests that high-impact loading is not necessarily linked to running injuries. For example, several studies have reported that knee osteoarthritis was found in equal frequency in runners and nonrunners [6] despite the fact that runners would incur more foot/ground contacts. Further, in a recent study, Miller [7] reported that joint loading in runners does not initiate knee osteoarthritis. Additionally, research has shown that joint contact forces in the lower limbs are 3 5 times smaller during the impact phase compared with the active phase of stance [8,9]. However, overuse injuries during running do not typically occur during the active phase of stance. As force is associated with acceleration, one would assume that runners who run faster and, therefore, increase their GRFs and would have higher incidence of injury; however, this is not evident in the literature. In fact, some level of force is necessary for bone health. It appears that, unlike previous thinking, high impact peaks nor high loading rates do not relate to injury [10]. However, in the biomechanics literature, there are still several citations that support the relationship between vertical GRF parameters and injury [11].

38.3.2 Hindfoot stability The motion of the hindfoot is a naturally occurring movement of the foot. It involves the tri-planar movements of hindfoot eversion, dorsiflexion, and forefoot external rotation. While “pronation” is the commonly used term (see elsewhere in this book) , it is actually calcaneal (hindfoot) eversion that is considered a major factor in the etiology of lower extremity soft tissue injury. There has been a substantial amount of research on the injury risk associated with hindfoot instability with particular focus on the measurement of hindfoot eversion. Runners with “excessive hindfoot eversion” have been classified as such by static and/or dynamic movement of the subtalar joint axis with reference to the lower leg. Interestingly, while commonly used, there is no definition for what is “excessive” regarding hindfoot eversion. The theory of hindfoot instability as a risk factor for injury involved the coupling and timing of the lower-extremity segments and joints. When calcaneal eversion occurs, there is simultaneous tibial internal rotation and knee flexion during impact phase of support. These actions are reversed during the push-off phase of support with hindfoot inversion, plantarflexion, and forefoot internal rotation. It has been postulated that too much hindfoot eversion increases the risk of injury by the asynchronous timing of maximum of tibial internal rotation and knee flexion causing increased stress to the knee [12]. These assumptions have not been verified. Shoe motion control features include medial arch support (anteriorly or posteriorly placed) and medial posting to prevent “excessive” eversion during midstance. Studies have assessed the effectiveness of motion control orthoses on hindfoot motion utilizing intra-cortical pins [13] and motion control footwear in three-dimensional (3D) biomechanical assessments of running. These studies have shown that considerable midsole variations resulted in very little difference in the maximum amount of eversion (Fig. 38.5).

FIGURE 38.5 Hindfoot eversion in A, B and C footwear. These footwear had different midsole designs. Note the less than one degree difference in maximum eversion angle.

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Athletic footwear research is typically assessed on healthy runners and, as results vary from individual to individual; perhaps we should identify functional groups of runners who may respond or not respond to particular footwear interventions. Functional groups have been suggested by the research group at the Human Performance Lab, University of Calgary [14]. These researchers suggested that, by using big data and utilizing machine learning techniques such as support vector machines, we may be able to identify these groups, who may be as small as one runner and extend to large groups of runners. Classifying runners into functional groups (i.e., a group of individuals who react/respond to an intervention in a similar way) may allow more accurate prescription of footwear for individual runners.

38.4

Types of shoes and their features

Running shoe design is focused on the principles of cushioning (ability to absorb shock) and stability (controlling hindfoot eversion). Other factors such as traction (friction with the surface) and durability (ability to withstand repetitive load, that is, multiple foot strikes) have also been considered but to a much lesser extent. Footwear design has aimed to specifically target these functions with biomechanical investigations.

38.4.1 Casual shoes Casual shoes came into fashion following World War 1. Following the baby boom in the 1950s, the sales of casual shoes increased again as school dress codes became more relaxed and these footwear became popular among the American youth. Early casual footwear designs consisted of solid rubber soles with a canvas upper, similar to Converse All Stars seen on the shelf today. Sales of sneakers (i.e., athletic type shoes) soared to 600 million pairs a year in 1957, which increased again in the 1970s when jogging became a popular activity. Subsequently, sneakers were developed not only for casual wear but for various sporting and lifestyle activities. The construction of casual sneakers involves often a zero heel-toe drop (i.e., the heel and the forefoot are at the same level) and flexible midsoles. These features ensure there is minimal antagonistic activity of the foot against the shoe.

38.4.2 Running shoes Compared to casual walking shoes, running and exercise footwear have received the most attention in terms of biomechanics of footwear design. Generally, exercise shoes have greater material added into the midsole. In particular, most exercise shoes have an increase in heel flare and medial posting to counter higher ranges of motion and forces experienced in the lower extremity. Medial posting and varus wedges are typically built into trail type footwear as hikers may experience greater hindfoot eversion on uneven surfaces. Much attention has been placed on controlling the hindfoot although many different types of shoe design have similar levels of calcaneal eversion (Fig. 38.5). Running shoes have several features to target a runner’s activity, anatomy, physiology, and kinematics. To accommodate many different types of runners, footwear manufacturers have categorized running footwear into three types of shoes based on the lower extremity characteristics of the runner. These categories are the following: (1) neutral; (2) motion control; and (3) stability. The majority of runners wear neutral (i.e., footwear that have little or no features that control the foot during foot/ ground contact) and motion control shoes. These shoes have a heel-toe drop of 10 12 mm on average, provide varying degrees of cushioning, and have arch support. Stability and motion control shoes differ in that they provide somewhat different levels of medio-lateral control of the foot. These shoes have features that typically consist of stiffer cushioning, stiffer heel counters, medial posts, and varus wedges. These features can be incorporated into the shoe to either stiffen the midsole and the upper to restrain movement of the foot and, therefore, limit hindfoot eversion, or they can modify the geometry of the cushioning to reduce the lever arm of the GRF [15] and thus reduce the eversion moment. A combination of both of these features where shoes are constructed with a cushioned and raised lateral midsole and firm medial posting have been shown to control hindfoot eversion [16].

38.4.3 Racing flats & spikes Racing flats and spiked shoes have several differences in their construction compared to typical running shoes. These include minimal heel-toe drop (also referred to as “zero drop”), reduced cushioning, and reduced shoe mass. Frederick

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et al., [17] has shown that metabolic cost is reduced by 1% for every 100 gram reduction in shoe mass. However, this was for running speeds under 3.5 m/s and may not be the same for competitive running. Loading rates, braking impulse, and peak vertical GRFs have been shown to be higher in racing flats and spikes compared with typical running shoes, likely related with reductions in stance time [18]. Spikes differ from racing flats only in that they house a spike plate in the forefoot of the outsole. Dependent on the race, the spike plate will differ in profile, number, and flexibility with shorter races requiring stiffer (often plastic vs rubber), more, and longer spikes.

38.4.4 Marathon shoes Over the course of a marathon race, a runner will experience around 25,000 foot-ground impacts. It has been reported that 74% 94% of marathon finishers adopt a hindfoot strike pattern during a race [14,19]. This highlights the importance of maintaining cushioning and stability features in marathon running shoes. Performance aspects have not been factored into footwear design until recently. Nike initiated their “Breaking2” project where, through a series of environmental constraints, they aimed to attempt a sub-2-hour marathon attempt. At the time of writing, the current world record is 2:02:57 set by Dennis Kimetto (Kenya) at the Berlin marathon in 2014. A sub-2-hour marathon requires a 2.5% improvement than this current record. As such, Nike aimed to breach this 2.5% threshold by optimizing drafting, course design, and running shoes [20]. Body weight support and forward propulsion during the stance phase contribute up to 80% of the metabolic cost of running [21], and therefore it was hypothesized that this can be modified by the foot-shoe-ground interface. The first method by which a runner may be able to reduce metabolic cost is by reducing the mass of the shoe, however, recent literature proposes that it is important to still maintain cushioning of the shoe, otherwise this will increase metabolic cost [17,22]. Franz et al. [22] compared running in lightweight racing flats (150 g) and barefoot running and found via extrapolation of their data that slightly lighter shoes (130 g) with the same cushioning properties would provide a 2.5% advantage over barefoot running. It is, therefore, advantageous for runners to wear a lightweight shoe with cushioning properties. The material of the cushioning is important since lighter weight resilient cushioning materials are desirable. Midsole plates can also enhance running economy by increasing longitudinal bending stiffness of the shoe. Research has shown that by adding a carbon fiber plate throughout the length of the midsole (longitudinal bending stiffness 5 38 N/mm) increased running economy by 0.8% compared with the same shoe with no plate (longitudinal bending stiffness 5 18 N/mm) [23]. With all of these features in mind, Nike designed the “VaporFly.” This B185 g shoe is constructed with lightweight ZoomX Pebax midsole foam and has a thin, stiff 100% carbon fiber plate with a slight forward inclination tuned to each individual athlete and a 10 mm heel-toe drop with a single breathable upper. Despite the shoe technology in combination with drafting, pacing, and course design, Eliud Kipchoge, missed the two hour barrier by 25 seconds, beating the record time by 2 minutes and 32 seconds (2.1% improvement) and his own personal record by 2 minutes and 40 seconds (2.2% improvement). Two independent studies have shown the Nike VaporFly (NVF) improved running economy by over 4% compared with marathon shoes (e.g., Nike Zoom and Adidas Adizero Adios) in controlled laboratory experiments [24,25]. These metabolic savings were explained in a follow up biomechanics study by Hoogkamer and colleagues . The authors conclude that the metabolic savings found in the NVF shoes may be attributed to: (1) better energy return in the new midsole foam; (2) the carbon fiber plate changing the lever arm at the ankle joint which reduced ankle extensor moments, and; (3) the stiffening of the metatarsophalangeal (MTP) joint caused by the carbon fiber plate which reduced dorsiflexion and negative work at the MTP joint [26].

38.4.5 Other sports shoes Footwear design considerations for team sports such as basketball and soccer must factor in the sudden accelerations, decelerations, changes of direction, jumping, and landing maneuvers which characterize these sports. Consequently, basketball and soccer shoes incorporate additional features compared to running shoes, often coming at an additional weight cost (Fig. 38.6). Cleats in the outsole of soccer shoes have been added to improve traction on grass and artificial turf surfaces. Initially, factory work boots were worn to play soccer, with some players adding metal studs. Soccer boots were not developed until the 1800s and were heavy (500 g), with leather studs and lacing systems still extending above the ankle [27]. It was not until the 1960s that below the ankle soccer shoes were created and used widely. Many comparisons have been made between bladed and studded cleat designs in soccer shoes for performance (i.e., ability to provide

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FIGURE 38.6 Footwear characteristics of running versus basketball and soccer shoes.

traction during change of direction maneuvers) [28,29] and minimizing injury risk potential (i.e., reducing peak knee and ankle joint moments) [30 32]. Research has shown that, compared to running shoes, soccer cleats increase rotational traction resulting in increased knee and ankle transverse and frontal plane joint moments during turning movements [31]. A configuration with minimal number of cleats and decreased cleat size is suggested to be beneficial to reducing injury risk while maintaining performance [27]. Impact forces during landing can reach up to 14 times body weight compared to 2 3 bodyweights during running [33]. Consequently, basketball shoes have been designed to have enhanced cushioning to attenuate these forces. Additionally, due to the high incidence of ankle and knee ligament sprain injuries in basketball [34], the upper of basketball shoes extends up over the ankle joint. High-top basketball shoes have been shown to have mixed success in reducing ankle and knee joint excursions and moments and have mixed perceptions of comfort [35]. Consequently, the most recent basketball shoes tend to have slightly lower uppers around the ankle, but not as low as running shoes.

38.4.6 New shoe innovations 38.4.6.1 Footwear embedded energy harvester InStep NanoPower, Vibram, Instituto Italiano di Tecnologia and University of Wisconsin-Madison have developed an in-shoe system that harvests the energy produced from walking. The insert converts mechanical energy to electricity via a microfluidic device where thousands of tiny droplets of liquid interact with a nanostructured substrate. The device has a power density of up to one kilowatt per 1.0 m2, with a capacity of 20 watts generated and stored in a rechargeable battery that takes four hours of walking to charge. Attached to the “Instep Nanopower” insert is an adapter into which wearers can plug their phones or other mobile devices.

38.4.6.2 Lacing systems Few biomechanical studies have investigated the lacing systems in athletic shoes. Hagen and Hennig [36] investigated the mechanical coupling of the foot and shoe on the protective properties of the shoe as it pertained to the lacing of the shoe. They suggested that a badly laced running shoe may allow the foot to slip within the shoe thus reducing the benefit of the design features of the shoe to the runner. These researchers undertook a study to explore the effect of shoe lacing on the biomechanics or runners. They reported that a firm foot-to-shoe coupling with an effective lacing system

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would lead to a more effective use of running shoe features and is likely to reduce the risk of lower limb injury. Consequently, several lacing innovations have been designed: G

G

Boa Lacing, now utilized by more than 300 brands, has stainless steel wires wrapped in low-friction lace guides in place of standard shoe lacing. These wires are connected to a dial that increases tension on the laces when twisted allowing the wearer to quickly and easily find the exact fit they need at any given time. The dial locks such that the laces are always firmly kept in place. Nike Hyper Adapt self-lacing shoes use a lacing system that is linked to sensors placed in the midsole of the shoe. The midsole senses a runner’s body mass and then feeds this information to a motor that tightens the laces based on this feedback. Adjustments can be made during wear via buttons placed on the upper. Nike Flywire cabling technology is a lacing system that is woven under the arch and through the upper of the shoe to provide a snug fit around the midfoot.

38.4.6.3 3D printed midsoles and outsoles Since 2013, several footwear companies have started 3D printing their midsoles. Fit-station powered by HewlettPackard is a new platform that delivers custom fit footwear through a combination of 3D foot scanning, gait analysis, and Multi Jet Fusion 3D printing technology. The foot scanning technology utilizes nine cameras to calculate foot length, width, and arch height, and recommendations can be made on features such as cushioning or placement of the arch. This analysis is also used to provide off-the-shelf shoe recommendations, custom 3D printed insoles, and polyurethane injected midsoles. These midsoles provide specific hardness zones and volumetric fit built to individual requirements (e.g., medial posting). Nike have also developed personalized football cleats to maximize agility. The shapes of the cleats are designed for sudden changes of direction and have a groove in the forefoot to assist with toe-off. These cleats can be individualized to player position such as in American football. For example, if the player typically changes direction on their left leg due to being on the left side of the field, the cleats on the left shoe will differ from the right.

38.4.7 Graphene outsoles Graphene, an atom-thin material 200 times stronger than steel, has recently been incorporated into the outsole of athletic footwear. This technology is expected to improve the strength while maintaining the flexibility and weight of the shoe and is expected to improve heat dissipation.

38.5

Shod versus barefoot

In the last several years, barefoot running and walking have been proposed as healthy alternatives to shod running/walking [37,38]. The most consistent findings when comparing barefoot to shod running are a reduction in stride length and a less dorsiflexed ankle at initial contact [37]. Other findings include a reduction in knee flexion excursion and knee joint moments during stance [39]. These studies have shown that barefoot running, in particular, cause runners to change their footfall strike from a hindfoot, heel-toe pattern, to one in which the point of the foot-ground collision is more anterior on the foot (i.e., on the metatarsal heads). These more anterior patterns are referred to as a midfoot strike in which the initial contact is on the metatarsal heads with the heel subsequently contacting the ground and the forefoot strike in which the initial contact is the same as the midfoot strike but the heel does not touch the ground. The majority of runners (i.e., 75% 80%) of runners employ a hindfoot strike with most of the remainder using midfoot strike [40]. Very few runners are habitual forefoot strikers. A forefoot strike pattern has been suggested to lower the risk of injury by reducing impact forces [38], but this has not been demonstrated in the literature. However, research has shown that, when habitual hindfoot runners transition to a forefoot strike, they have higher energy cost [41], overload untrained tissues [42], and consequently have higher incidence of injury to the foot and ankle region [43]. In spite of this evidence, many coaches and trainers suggest that runners would benefit by altering their footfall pattern from a hindfoot to a more anterior midfoot or forefoot pattern. They suggest that such a change would decrease the foot-ground impact and reduce running-related injuries. Hamill and Gruber [19] suggested that neither of these supposed “benefits” would occur. Overall, there does not seem to be compelling evidence for either barefoot running or changing a footfall pattern from a hindfoot to a midfoot or forefoot pattern.

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Footwear related injuries

There is much debate in the literature regarding whether or not athletic footwear influence running injury rates both positively and negatively. For the ever-growing population of runners, 19% 79% experience running related injuries every year, with the majority of these injuries being to the knee joint (Fig. 38.7) [44 46]. Previous injury, training errors, and poor running mechanics have been associated with these injuries [46]. Despite the development and innovation of athletic footwear since the 1970s, running injury incidence remains unchanged [46]. The lack of change in injury incidence may be due to several factors: (1) the changing running population; whereas runners in the 1970s and 1980s were competitive runners and the majority were males, now most runners are primarily recreational with a slight majority are female [45,47]; (2) the variation of how running injuries are defined between studies [47,48]; and (3) the conflicting evidence for previously researched mechanisms (i.e., impact forces and hindfoot stability) related to footwear and injury. To appropriately design footwear to assist in preventing injury, it is first important to understand the mechanisms and countermeasures of running related injuries. The most robust research design is to perform prospective studies wherein runners are assessed biomechanically and then followed for a period of time. Over time, the injuries to runners would be cataloged and related to factors such as footwear usage, training history and age. Unfortunately, due to the labor involved in collecting biomechanical data in large populations, most prospective studies do not include biomechanics of groups or types of runners (i.e., elite/novice, male/female, high arch/low arch, etc.). Additionally, it is difficult to compare studies in which there are many ways of defining injury: that is, injuries per 1000 km [52], injuries per 100 runners [46], and number of injured runners per 1000 hours of running [53]. As to impact forces, a systematic review and meta analysis [10] has shown no difference in impact peak or active peak of the vertical GRF component in injured and noninjured runners (n 5 1000). There has been some evidence for the link between loading rate, “excessive” hindfoot eversion, and medial tibial stress syndrome [18,54]. Messier and colleagues [45] reported that, in a prospective 2-year follow up of runners, the lone predictor of injury was increased knee stiffness. Parameters such as flexibility, arch height, Q-angle, hindfoot eversion, strength, mileage, footwear, and previous injury were not found to be etiologic factors for overuse injury. While numerous studies have provided the research and the industry community with innovative footwear technologies such as gels, foams, midsole posting designs, and encapsulated gases, the goal of empirically reducing injuries has yet to be achieved. Numerous studies assigning footwear to foot type [55], utilizing shock absorbing insoles [56], using soft vs hard midsoles [5], and motion control footwear [57] have found no influence on the incidence of running-related injuries. However, some studies have shown that running-related injury risk was lower among runners using motion control shoes compared with those running with neutral footwear over a six-month period [58]. With respect to minimalist shoes, conflicting research has shown both increases in injury incidence and lower limb pain [59] and 3.4 times less injury incidence than shod runners [60]. However, the difference between these studies is that runners were either experienced or not experienced in running in minimalist footwear and/or were habitual forefoot runners.

FIGURE 38.7 Lower limb injury site as a percentage of total running injury incidence from Clement et al. [49], MacIntyre et al. [50], Taunton et al. [51] and Messier et al. [46,50,51,53][45].

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38.7

Future footwear research

Biomechanics has had a significant impact on the design of footwear. With the conflicting and contrasting evidence for the biomechanical factors associated with running injury, footwear design parameters, and the influence of these parameters on both biomechanics and injury incidence, several new paradigms have been suggested. That is, the methods that have been used to evaluate the design parameters of footwear may need to be changed. Of particular interest, Professor Benno Nigg at the University of Calgary suggested that we re-think our notions of cushioning and hindfoot stability as risk factors for running-related injuries. Nigg suggested an alternative paradigm that he termed as the “preferred movement path” [61]. He suggested that there is an individual- and task-specific locomotion pattern that is determined by many factors (i.e., muscles, tendon, ligaments, bone structure, etc.), and that this preference for a particular motion path may explain why shoes have little or only a moderate effect on lower extremity kinematics. This new paradigm may lead to new assessment tools such as how footwear and may affect the determination of a runner’s deviation from the runner’s preferred path. It is obviously very difficult to ascertain a runner’s preferred movement path. Rather, it may be easier to determine their habitual path or the path that they use most of the time. One particular footwear company has conducted research to develop a paradigm related to the choice of footwear to their habitual path. Researchers at this company determined that the knee should be the focus of the paradigm because the largest number of injuries are to the knee [61]. A method was determined to calculate if a person deviated from their habitual path and how this deviation could be affected by a footwear change. For example, if an individual had a high deviation from their habitual path, it might be beneficial to this runner to use a particular shoe that can stabilize their knee movement. On the other hand, a low deviator may need nothing more than a neutral shoe to maintain their habitual motion path. There are other directions that could be taken to improve/evaluate running footwear designs that also show promise. The concept of the “eversion buffer” [62] or the prolonged loading calculating bone stress [63] illustrate differences in injury between groups and between shoes. The “eversion buffer” concept hinges on the proximity of the calcaneal eversion angle to the limit of eversion available to the individual. The closer the proximity the more likely to incur an eversion-like injury. Thus, footwear would be designed with features to bring the individual’s eversion angle towards the middle of the range of eversion motion. Calculating bone stress during a prolonged run may also be useful in the design of footwear. Other suggestions include using sensors in the shoe to evaluate footwear during both steady state and fatigued conditions. Sensors in the shoe could result in large quantities of data creating population statistics on running parameters in multiple environments. Great strides have also been made in the development of analysis techniques of evaluating footwear. Techniques such as principal components analysis, functional data analysis, wavelets, and dynamical system analysis can be used with large quantities of data to design and evaluate the effectiveness of footwear. In conclusion, it is difficult to determine the effectiveness of footwear in either decreasing the injury risk or improving the performance of individuals in athletic contests or in daily activities. It is clear that biomechanics has had a significant impact on the development of both athletic and casual footwear. However, the biomechanics of human movement and the interaction of the individual and the shoe must continue to be evaluated using appropriate biomechanical analyses.

References [1] [2] [3] [4] [5] [6] [7] [8] [9] [10]

Cheskin M, Sheskin KJ, Bates BT. The complete handbook of athletic footwear. Fairchild Publ 1987;. Cavanagh PR. The running shoe book. Anderson World 1980;. Cressman LS. Western prehistory in the light of carbon 14 dating. Southwest J Anthropol 1951;7:289 313. Radin EL, Parker HG, Pugh JW, Steinberg RS, Paul IL, Rose RM. Response of joints to impact loading—III: relationship between trabecular microfractures and cartilage degeneration. J Biomech 1973;6:51 7. Hardin EC, Van Den Bogert AJ, Hamill J. Kinematic adaptations during running: effects of footwear, surface, and duration. Med Sci Sports Exerc 2004;36:838 44. Hansen P, English M, Willick SE. Does running cause osteoarthritis in the hip or knee? PM&R 2012;4:S117 21. Miller RH. Joint loading in runners does not initiate knee osteoarthritis. Exerc Sport Sci Rev 2017;45:87 95. Cole GK, Nigg BM, Fick GH, Morlock MM. Internal loading of the foot and ankle during impact in running. J Appl Biomech 1995;11:25 46. Scott SH, Winter DA. Internal forces of chronic running injury sites. Med Sci Sports Exerc 1990;22:357 69. van der Worp H, Vrielink JW, Bredeweg SW. Do runners who suffer injuries have higher vertical ground reaction forces than those who remain injury-free? A systematic review and meta-analysis. Br J Sports Med 2016;8:450 7.

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[11] Milner CE, Ferber R, Pollard CD, Hamill J, Davis IS. Biomechanical factors associated with tibial stress fracture in female runners. Med Sci Sports Exerc 2006;38:323 8. [12] Hamill J, Bates BT, Holt KG. Timing of lower extremity joint actions during treadmill running. Med Sci Sports Exerc 1992;24:807 13. [13] Arndt A, Wolf P, Liu A, Nester C, Stacoff A, Jones R, et al. Intrinsic foot kinematics measured in vivo during the stance phase of slow running. J Biomech 2007;40:2672 8. [14] Hoerzer S, von Tscharner V, Jacob C, Nigg BM. Defining functional groups based on running kinematics using self-organizing maps and support vector machines. J Biomech 2015;48:2072 9. [15] Shorten MR. Running shoe design: protection and performance. Marathon Med 2000;159 69. [16] Nigg BM, Bahlsen HA. Influence of heel flare and midsole construction on pronation supination and impact forces for heel-toe running. Int J Sport Biomech 1988;4:205 19. [17] Frederick EC, Daniels JR, Hayes JW. The effect of shoe weight on the aerobic demands of running. In: Bachl N, Prokop L, Suckert, R, editors Current topics in sports medicine, proceedings of the world congress of sports medicine. Urban and Schwarzenberg, Vienna; 1984. p. 616 625. [18] Logan S, Hunter I, Hopkins JT, Feland JB, Parcell AC. Ground reaction force differences between running shoes, racing flats, and distance spikes in runners. J Sports Sci Med 2010;9:147. [19] Hamill J, Gruber AH. Is changing footstrike pattern beneficial to runners? J Sport Health Sci 2017;6:146 53. [20] Hoogkamer W, Kram R, Arellano CJ. How biomechanical improvements in running economy could break the 2-hour marathon barrier. Sports Med 2017;47:1739 50. [21] Arellano CJ, Kram R. Partitioning the metabolic cost of human running: a task-by-task approach. Oxford University Press; 2014. p. 1084 98. [22] Tung KD, Franz JR, Kram R. A test of the metabolic cost of cushioning hypothesis during unshod and shod running. Med Sci Sports Exerc 2014;46:324 9. [23] Roy J-PR, Stefanyshyn DJ. Shoe midsole longitudinal bending stiffness and running economy, joint energy, and EMG. Med Sci Sports Exerc 2006;38:562 9. [24] Barnes KR, Kilding AE. A randomized crossover study investigating the running economy of highly-trained male and female distance runners in marathon racing shoes vs track spikes. Sports Med 2018;1 12. [25] Hoogkamer W, Kipp S, Frank JH, Farina EM, Luo G, Kram R. A Comparison of the Energetic Cost of Running in Marathon Racing Shoes. Sports Med 2018;48:1009 19. [26] Hoogkamer W, Kipp S, Kram R. The biomechanics of competitive male runners in three marathon racing shoes: a randomized crossover study. Sports Med 2018;1 11. [27] Hennig EM, Sterzing T. The influence of soccer shoe design on playing performance: a series of biomechanical studies. Footwear Sci 2010;2:3 11. [28] Andreasson G, Lindenberger U, Renstro¨m P, Peterson L. Torque developed at simulated sliding between sport shoes and an artificial turf. Am J Sports Med 1986;14:225 30. [29] Shorten M, Hudson B, Himmelsbach J. Shoe-surface traction of conventional and in-filled synthetic turf football surfaces. XIX Int Congr Biomech 2003;. [30] Queen RM, Charnock BL, Garrett WE, Hardaker WM, Sims EL, Moorman CT. A comparison of cleat types during two football-specific tasks on FieldTurf. Br J Sports Med 2008;42:278 84. [31] Stefanyshyn DJ, Lee J-S, Park S-K. The influence of soccer cleat design on resultant joint moments. Footwear Sci 2010;2:13 19. [32] Torg JS, Quedenfeld TC, Landau S. The shoe-surface interface and its relationship to football knee injuries. J Sports Med 1974;2:261 9. [33] Panzer VP, Wood GA, Bates BT, Mason BR. Lower extremity loads in landings of elite gymnasts. Biomech XI-B 1988;727 35. [34] Hootman JM, Dick R, Agel J. Epidemiology of collegiate injuries for 15 sports: summary and recommendations for injury prevention initiatives. J Athl Train 2007;42:311. [35] Verhagen EA, van der Beek AJ, van Mechelen W. The effect of tape, braces and shoes on ankle range of motion. Sports Med 2001;31:667 77. [36] Hagen M, Hennig EM. Effects of different shoe-lacing patterns on the biomechanics of running shoes. J Sports Sci 2009;27:267 75. [37] Altman AR, Davis IS. Barefoot running: biomechanics and implications for running injuries. Curr Sports Med Rep 2012;11:244 50. [38] Lieberman DE, Venkadesan M, Werbel WA, Daoud AI, D’andrea S, Davis IS, et al. Foot strike patterns and collision forces in habitually barefoot vs shod runners. Nature 2010;463:531. [39] Sinclair J. Effects of barefoot and barefoot inspired footwear on knee and ankle loading during running. Clin Biomech 2014;29:395 9. [40] Hasegawa H, Yamauchi T, Kraemer WJ. Foot strike patterns of runners at the 15-km point during an elite-level half marathon. J Strength Cond Res 2007;21:888. [41] Gruber AH, Umberger BR, Braun B, Hamill J. Economy and rate of carbohydrate oxidation during running with rearfoot and forefoot strike patterns. J Appl Physiol 2013;115:194 201. [42] Stearne SM, Alderson JA, Green BA, Donnelly CJ, Rubenson J. Joint kinetics in rearfoot vs forefoot running: implications of switching technique. Med Sci Sports Exerc 2014;46:1578 87. [43] Ridge ST, Johnson AW, Mitchell UH, Hunter I, Robinson E, Rich BS, et al. Foot bone marrow edema after a 10-wk transition to minimalist running shoes. 2013;45:1363 1368. [44] van Gent BR, Siem DD, van Middelkoop M, van Os TA, Bierma-Zeinstra SS, Koes BB. Incidence and determinants of lower extremity running injuries in long distance runners: a systematic review. Br J Sports Med 2007;469 80.

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[45] Messier SP, Martin DF, Mihalko SL, Ip E, DeVita P, Cannon DW, et al. A 2-year prospective cohort study of overuse running injuries: the runners and injury longitudinal study (TRAILS). Am J Sports Med 2018;46:2211 21. [46] Van Mechelen W. Running injuries. Sports Med 1992;14:320 35. [47] Videbaek S, Bueno AM, Nielsen RO, Rasmussen S. Incidence of running-related injuries per 1000 h of running in different types of runners: a systematic review and meta-analysis. Sports Med 2015;45:1017 26. [48] Bahr R. No injuries, but plenty of pain? On the methodology for recording overuse symptoms in sports. Br J Sports Med 2009;43:966 72. [49] Clement DB, Taunton JE. A guide to the prevention of running injuries. Australian family physician 1981;10(3):156 61. [50] Macintyre J, Taunton J, Clement D, et al. Running injuries: a clinical study of 4,173 cases. Clinical Journal of Sports Medicine 1991;1:81 7. [51] Taunton JE, Ryan MB, Clement DB, McKenzie DC, Lloyd-Smith DR, Zumbo BD. A retrospective case-control analysis of 2002 running injuries. British journal of sports medicine 2002;36(2):95 101. [52] Gerlach KE, Burton HW, Dorn JM, Leddy JJ, Horvath PJ. Fat intake and injury in female runners. J Int Soc Sports Nutr 2008;5:1. [53] Jakobsen BW, Kroner K, Schmidt SA, Kjeldsen A. Prevention of injuries in long-distance runners. Knee Surg Sports Traumatol Arthrosc 1994;2:245 9. [54] Zadpoor AA, Nikooyan AA. The relationship between lower-extremity stress fractures and the ground reaction force: a systematic review. Clin Biomech 2011;26:23 8. [55] Knapik JJ, Brosch LC, Venuto M, Swedler DI, Bullock SH, Gaines LS, et al. Effect on injuries of assigning shoes based on foot shape in air force basic training. Am J Prev Med 2010;38:S197 211. [56] Withnall R, Eastaugh J, Freemantle N. Do shock absorbing insoles in recruits undertaking high levels of physical activity reduce lower limb injury? A randomized controlled trial. J R Soc Med 2006;99:32 7. [57] Nielsen RO, Buist I, Parner ET, Nohr EA, Sørensen H, Lind M, et al. Predictors of running-related injuries among 930 novice runners: a 1-year prospective follow-up study. Orthop J Sports Med 2013;1:1 7. [58] Malisoux L, Chambon N, Delattre N, Gueguen N, Urhausen A, Theisen D. Injury risk in runners using standard or motion control shoes: a randomised controlled trial with participant and assessor blinding. Br J Sports Med 2016;0:1 7. [59] Ryan M, Elashi M, Newsham-West R, Taunton J. Examining injury risk and pain perception in runners using minimalist footwear. Br J Sports Med 2014;48:1257 62. [60] Goss DL, Gross MT. Relationships among self-reported shoe type, footstrike pattern, and injury incidence. US Army Med Dep J 2012;. [61] Nigg BM, Vienneau J, Smith AC, Trudeau MB, Mohr M, Nigg SR. The preferred movement path paradigm: influence of running shoes on joint movement. Med Sci Sports Exerc 2017;49:1641 8. [62] Rodrigues P, TenBroek T, Van Emmerik R, Hamill J. Evaluating runners with and without anterior knee pain using the time to contact the ankle joint complexes’ range of motion boundary. Gait Posture 2014;39:48 53. [63] Firminger C, Edwards B. Effects of minimalist footwear and stride length reduction on the probability of metatarsal stress fracture: a weibull analysis with bone repair. Footwear Sci 2017;9:S127 9.

Chapter 39

Minimal Shoes: Restoring Natural Running Mechanics Karsten Hollander1 and Irene S. Davis2 1

Spaulding National Running Center, Physical Medicine and Rehabilitation, Harvard Medical School, Cambridge, MA, United States, 2Spaulding

National Running Center, Physical Medicine and Rehabilitation, Harvard Medical School, Cambridge, MA, United States

Abstract The purpose of this chapter is to examine the benefits (and pitfalls) of minimal shoe running. We will begin with a brief history of the evolution of the modern running shoe from minimal to modern conventional shoes. We will then review the mechanics of barefoot running as it is our most natural running form and one that minimal footwear is aimed at promoting. We will then compare barefoot running mechanics to both conventional and minimal shoes. We will also discuss the benefits of minimal footwear on the musculoskeletal system. Proper preparation and transition to minimal footwear will be reviewed. Finally, we will conclude with areas of future research.

39.1

Introduction

Homo Sapiens have been running for nearly 2 million years, with the large majority of this time being barefoot. Therefore, the modern human should be very well adapted to barefoot locomotion [1,2]. The first shoes emerged approximately 10,000 years ago. They were constructed of sage brush bark and consisted of a foot bed and straps to hold it onto the foot (Fig. 39.1). They were very minimal by today’s standards and primarily served as protection for the sole of the foot. Athletic shoes eventually appeared nearly 200 years ago and again were very minimal. However, with time, athletic shoes, in particular running footwear, have become increasingly more cushioned and supportive. That is, until the renewed interest in barefoot running that accompanied the best-selling book Born to Run by Christopher McDougall [3]. This interest was a catalyst for the development of footwear that fostered mechanics similar to that of barefoot running. However, many simply replaced their mileage in conventional footwear with minimal shoes without proper physical preparation [4 6]. Unfortunately, these shoes with their reduced cushioning and support placed greater demands on the foot and lower leg and requires greater capacity of these muscles [7,8]. Without adequate transition, injuries began to be reported [4 6]. Rather than taking the time to build the capacity of the foot and lower leg and slowly increase mileage, many of these runners returned to their conventional footwear, before gaining the benefits the minimal shoes might offer. Therefore, the purpose of this chapter is to examine the benefits and pitfalls of minimal footwear. We will compare running mechanics between minimal and conventional footwear as well as their effects on the musculoskeletal system.

39.2

Brief history of running footwear

New manufacturing processes in the industrial era led to the invention of athletic shoes [9,10]. It began in 1832 when leather soles were replaced with rubber ones. Rubber was made stronger and more elastic with the process of vulcanization in 1944, thus further improving athletic shoes [11]. At the end of the 19th century, the company J.W. Foster and Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00035-4 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 39.1 First-documented footwear, made of sagebrush, found in the Fort Rock caves in Oregon. These shoes date back 10,000 years. Copyright University of Oregon Museum of Natural and Cultural History, photograph by Jack Lui.

Sons developed the first running shoe by adding spikes under the forefoot region of athletic shoes. The company Keds (a footwear brand of the tire manufacturer US Rubber Company) then invented the first “sneaker” in 1916 [10] (Fig. 39.2). Running shoes were further customized in 1926 by Adi Dassler [9]. The development of the modern running shoe was accelerated in the 1960s and 1970s, with the founding of footwear companies such as Onitsuka Tiger (Kobe, Japan), which later became Asics, and Nike (Beaverton, OR, USA). Adi Dassler had founded Adidas (Herzogenaurach, Germany) in 1949 and J.W. Foster and Sons became Reebok (Boston, MA, USA) in 1958 [9]. The company New Balance (Boston, MA, USA) invented the “Trackster” in 1960 and Onitsuka Tiger manufactured the first running shoe with a small cushioned heel in 1964. Nike also introduced other footwear aspects such as rubber waffle soles and breathable nylon uppers in their Cortez shoe. With the “running and fitness boom” in the 1970s, there was an explosion of untrained individuals taking up running, and the prevalence of running-related injuries increased. This led to the first studies on running-related injuries being published [12,13]. At this time, shoes were still highly minimal, incorporating little to no cushioning or support. As a result of consultation with some sports podiatrists, it was suspected that these injuries were related to runners not being adapted to the minimal nature of the footwear. Therefore, significant changes were introduced to the running shoe, resulting in new footwear categories, such as cushioned, and motion control shoes [9]. Cushioned shoes were designed to reduce impacts, while motion control shoes were intended to reduce excessive eversion. However, these improvements have not resulted in a reduction of the running injury prevalence over the past 40 years [14]. Depending on length of study and definition of injury, these injuries are still reported to be as high as 79% per year [15]. Over the last 10 years, a re-emergence of a less cushioned, less supportive, and highly flexible ‘minimal’ shoe has dominated the scientific debate [9,16] (Fig. 39.3). This debate was catalyzed by the publication of the best-selling book

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FIGURE 39.2 Early sports shoes. Keds sneakers that were fashioned from vulcanized rubber and a canvas upper.

FIGURE 39.3 (A) Example of a minimal shoe that lacks support, has no midsole, and (B) is highly flexible.

Born to Run by Chris McDougall in 2009 [3]. At approximately the same time, Lieberman et al. published a scientific paper on barefoot and minimal footwear running that has been cited nearly 1400 times since its publication in 2010 [17]. Both of these publications proposed possible advantages of either barefoot or minimal footwear running. They suggested this style of running would alter mechanics in such a way as to reduce vertical force impacts [18] that have been associated with running injuries [19 21]. To date, there has been some evidence regarding the ability of minimal shoes to reduce impacts [22] as well as injury rates [23,24]. However, too rapid a transition to minimally shod running has been associated with an increased risk for some injuries [4,5]. This has led many runners back to their conventional shoes.

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FIGURE 39.4 Phases of barefoot gait. (A) Plantarflexion at foot strike. (B) Inversion at foot strike. (C) Mild lowering of arch at midstance. (D) Inversion and raising of arch at push-off.

39.3

Biomechanics of barefoot and conventional shod running

39.3.1 Barefoot running pattern Humans evolved running barefoot; therefore, it can be considered as our most natural state [1]. As minimal running shoes were designed to mimic barefoot running, we will first describe the mechanics of barefoot running, using those who have grown up barefoot and therefore are likely the most habituated [25]. Most habituated adult barefoot runners land slightly plantarflexed (forefoot strike—FFS) on the lateral side of the ball of their foot with their arch raised and the hindfoot inverted [26] (Fig. 39.4A and B). Their foot lands beneath a flexed knee with a vertical tibia. The foot then simultaneously everts and dorsiflexes and the arch lowers slightly until the foot is flat at midstance (Fig. 39.5C), where the knee is maximally flexed [27]. During the last half of stance, the foot begins to invert and plantarflex as it propels the body forward. Push-off occurs with the arch raised and the ball of the foot flat on the ground, rolling forward until toe-off (Fig. 39.4D).

39.3.2 Conventional shod running pattern Runners habituated to conventional shoes typically land dorsiflexed and on the lateral side of their heel (rearfoot strike—RFS) with their hindfoot inverted [28]. (Editor’s note: Elsewhere in this book, we use “hindfoot” exclusively, but due to overwhelming convention in the existing literature, we will use “rearfoot” strike in this chapter.) Their leg is outstretched with their foot landing in front of their slightly flexed knee with a forwardly inclined tibia. The foot initially plantarflexes to achieve foot flat, and then simultaneously dorsiflexes and everts as the tibia moves forward achieving maximal knee flexion at midstance. The foot then plantarflexes and inverts during the last half of stance through push-off from the ball of the foot.

39.3.3 Comparison of mechanics between conventional shod and barefoot running One of the most important acute effect of the footstrike differences between shod and barefoot runners is on the ground reaction forces. A RFS pattern results in an abrupt vertical impact force as the heel strikes the ground, that is typically missing from an FFS pattern (Fig. 39.5A). The slope of the initial rise of the vertical force is referred to as the vertical load rate (Fig. 39.5B). Due to the abrupt nature of this rise in a RFS pattern, vertical load rate is significantly higher in a RFS than an FFS pattern [17,29]. An increased load rate have been associated with a variety of running-related injuries [19,20,30]. It is likely that it is discomfort or pain from this impact force that promotes an FFS pattern in barefoot runners. In fact, Wearing et al. [31] reported that the pressure pain threshold at the heel is exceeded when the loads applied to the heel exceed those of walking. This suggests that the heel pad is sufficient to prevent pain during walking, but not running. Another sensory benefit

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500 0

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FIGURE 39.5 (A) Typical vertical ground reaction force curve for a rearfoot striker (RFS) and a forefoot striker (FFS). (B) The loading rate of the vertical ground reaction force curve.

FIGURE 39.6 (A) Relationship between footwear history and footstrike angle. (B) Relationship between running history and footstrike angle. Note that those who are always barefoot and those who often run exhibit mostly an FFS pattern. Adapted from Lieberman et al., Plos One, 2015.

of landing on the ball of the foot is that it is naturally wider and has a larger surface area over which to distribute the ground reaction forces during running, thereby reducing the overall pressures experienced by the foot. Studies have confirmed the relationship between footwear and footstrike. Lieberman et al. [32] studied adolescent and adult runners from the Eldoret region of Kenya who run regularly as part of their daily lives. These authors reported that those who are always barefoot land on the ball of their feet, while those who are always shod land on their heels. Those who spend mixed time in shoes and barefoot tend to run with a midfoot strike pattern (Fig. 39.6A). However, this does not seem to be the case for habitual barefoot agropastoral individuals from the Hadzu region of Kenya who do not run regularly. When asked to run, these individuals predominantly use a rearfoot strike (RFS) pattern [33]. This is consistent with the findings of Lieberman that those who run often tend to use a FFS pattern (Fig. 39.6B). Results for children have been more varied. Hollander et al. [34] noted that 75% of 6 year old barefoot runners land with a RFS pattern, however by the time they are 14 years old, only 25% are rearfoot strikers. These findings are in contrast to those of habitually shod children [35 38] in which RFS patterns were found to be least prevalent in young children, but then increase with age. For example, Latorre et al. [38] reported that RFS patterns increased from only 40% in 3 4 year olds to 92% in 15 16 year olds. Therefore, over time, being habitually barefoot reduces RFS patterns, while being habitually shod increases them. As strike pattern is the primary difference between habituated barefoot and shod runners, the majority of the differences that are noted between them are at initial contact [39]. Both land with their foot in an inverted position. However, a barefoot runner typically lands in ankle plantarflexion and increased knee flexion [25,40,41]. In contrast, a shod runner typically lands in ankle dorsiflexion and greater knee flexion than a barefoot runner [25,40]. This results in a longer stride and lower cadence than the barefoot runner [17,42].

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FIGURE 39.7 Effect of the vertical ground reaction force (vGRF) on the eversion moment of the foot in the (A) barefoot condition and (B) shod condition. Note the longer moment arm in the shod condition due to the wider base in a conventional running shoe. Despite identical vGRFs, the longer moment arm contributes to a greater eversion moment in the shod condition.

Numerous studies have examined the barefoot condition in runners who are habitually shod. The novel barefoot condition often does not elicit a consistent FFS pattern. However, novice barefoot runners do typically exhibit a less dorsiflexed ankle angle at ground contact [43 45], a reduced rate of rearfoot strikes, and increased knee flexion at ground contact [40,45]. This results in an increased cadence with decreased step length and reduced contact time [40,43,45,46]. These findings are consistent for children and adolescent populations during novice barefoot conditions [47,48]. Furthermore, due to the narrower width of a bare foot compared to a shoe, the moment arm of the vertical reaction forces to the subtalar joint axis is smaller and results in reduced external eversion moments [45,49,50] (Figs. 39.7A and B). As a result, Bonacci et al. [45] has noted reduced hindfoot eversion in barefoot compared with shod running. Barefoot running has also been associated with decreased power absorption at the knee and increased power absorption at the ankle [22,40,46]. There have also been studies examining the effect of transitioning from shod to barefoot running. Hollander et al. [29] conducted a randomized control trial (RCT) that incorporated eight 15-minute weekly sessions of uninstructed barefoot running. This led to a less dorsiflexed foot when compared to the shod habituation group. However, this did not result in a reduction of overall loading rates [29]. This may be due to the relatively short amount of time for the barefoot habituation. Hashish et al. [51] also reported on a progressive 8 10 week uninstructed barefoot transition program that led to participants running 100% of their base mileage barefoot. However, there were no overall changes in lower extremity mechanics, likely due to the fact that only 3 of the 18 participants transitioned to a forefoot strike pattern. Tam et al. [52] conducted an 8-week, but uninstructed, barefoot training program where runners were progressed to 40 minutes of barefoot running. Again, this led to no overall biomechanical changes in these runners. As a result, participants did not change lower limb kinematics (ankle, foot, and knee angles at ground contact), flight time, contact time, or cadence. However, these authors analyzed a subset of individuals who significantly decreased their load rates. They found that these runners transitioned from a RFS pattern to an FFS pattern, likely resulting in the reduced load rates. One of the issues with these transition studies is that it is still unknown as to when one is habituated to a condition. It is very possible that more than 2 3 months is needed before a runner becomes fully habituated to a barefoot condition. Suggestions based upon length of time barefoot [53], percentage of mileage run barefoot [17,32,54], and percentage of overall barefoot activities have all been used [55 57]. What is clear from these relatively short transition studies is that instruction may be helpful in promoting changes associated with natural barefoot running.

39.4

Minimal footwear running

39.4.1 Definition of full minimalist footwear Prior to the 1970s, all running shoes were essentially minimal. However, since the emergence of the “Nike free” [58], all shoes with less cushioning and more flexibility have been marketed as minimalist shoes [59]. However, not all of the shoes in this category have the same potential to mimic barefoot running biomechanics [43]. In 2015, an effort was made to standardize the definition and in this consensus statement by Esculier et al. [60] a minimalist shoe is defined as

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“footwear providing minimal interference with the natural movement of the foot due to its high flexibility, low heel to toe drop, weight and stack height, and the absence of motion control and stability devices.” Furthermore, a “minimalist index” was developed that allocates points in five categories (weight, flexibility, heel to toe drop, stack height, and motion control/stability devices) [60]. This minimalist index ranges from 0% to 100% with 100% representing the highest degree for “minimalism.” For the purpose of this chapter, we will classify shoes as full minimal shoes (MI . 75%), partial minimal shoes (MI # 75% and $ 25%), and conventional running shoes (MI , 25%) (Fig. 39.8). The full minimal footwear consists of a low weight, no midsole, as well as a high flexibility and no motion control or stability elements. Of all footwear, the full minimal footwear has the highest potential to simulate barefoot running [43]. It provides more freedom of motion of the foot while offering protection of the sole of the foot from the elements, such as cold and heat, as well as sharp objects. However, it is not identical to barefoot running as some of the sensory input from the foot is lost. Numerous studies have shown that this sensory input is important to overall posture and gait [61 63].

39.4.2 Comparison of full minimal to barefoot running There have been several studies comparing full minimal footwear and barefoot running. A study by Squadrone and Gallozzi [44] showed full minimal footwear (Vibram five finger) to have the same contact times, foot strike index, and loading rate, as well as ankle, knee, and foot kinematics as barefoot running. Hollander et al. [43] found that running in full minimal footwear has the highest potential to mimic biomechanics of barefoot running but still differ in the ankle angle at ground contact, cadence, and step length. In another study by Squadrone et al. [64] several minimal footwear models were compared to barefoot and cushioned running shoe running. It was reported that less mass, heel stack height, and heel-toe drop in shoes are associated with the potential to simulate barefoot running when looking at foot strike index, foot angle at contact, and spatiotemporal parameters (cadence, contact time, and stride length).

39.4.3 Comparison of full minimal to partial minimal shoes A few studies have compared minimal to partial minimal shoes [43,45,64]. The most important characteristic that differentiates a partial minimal shoe to a minimal one is the cushioning, as it promotes the use of a rearfoot strike. Bonacci et al. [45] compared a partial minimalist shoe (Nike free) to barefoot running after providing a 10-day familiarization period. They reported significant differences in stride length, cadence, ankle angle at contact, peak knee flexion during stance, and ankle plantarflexion angle during toe-off as well as in knee (external rotation, abduction) and ankle (plantarflexion, inversion) moments, joint power, and joint work at the ankle and knee. When comparing several full and partial minimal shoes to each other, the full minimal shoes were closer to barefoot running and the partial minimal shoes were closer to cushioned footwear running when looking at foot strike parameters, cadence, contact time, and stride length [64]. Another study by Hollander et al. [43] compared running in a full minimal shoe (Leguano) to a partial minimal shoe (Nike Free), but without a familiarization period. They reported that the minimal shoe resulted in less dorsiflexion at ground contact, increased cadence, and reduced step length when compared to the partial minimal shoe. The rate of rearfoot strikes was 62% 74% in the minimal shoe and 88% 93% in the partial minimal shoes. Landing with a RFS pattern in a partial minimal shoe has also been reported by others [65]. This can lead to high loading rates due to the reduced cushioning, which might increase the risk for running-related injuries. In fact, Ryan et al. [66] reported twice the incidence of running-related injuries when running in a partial minimal compared to a minimal. These studies all suggest that, despite perceptions, partial minimal and minimal shoes are not alike. The primary

FIGURE 39.8 Examples of running shoes: (A) minimal, (B) partial minimal, (C) conventional.

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motivation to use this type of footwear has been reported to be to prevent running-related injuries [67]. As such, minimal shoes may better promote the forefoot strike pattern that is typically associated with lower impacts, which have been associated with decreased injuries [19 21,68].

39.4.4 Comparison of full minimal to conventional footwear The comparison of between running in full minimal to conventional running shoes is similar as when comparing barefoot and conventionally shod running. A runner in minimal footwear lands with a less dorsiflexed ankle angle and a reduced rate of rearfoot strikes, as well as a higher cadence and reduced step length [43,44]. Additionally, there is the presence of a vertical impact peak of the ground reaction force in the conventional shoe that is typically reduced or absent (depending on footstrike pattern) in the minimal shoe, similar to barefoot running (Fig. 39.4A) [40,47]. As a result, running with a RFS pattern in a conventional shoe typically results in greater vertical load rates than using an FFS pattern in a minimal shoe, and these loading rates have been associated with some running related injuries. Approximately 1% of runners in conventional shoes naturally land on their forefoot [69]. While these runners have reduced vertical loading rates compared to those with conventional RFS, and they have increased mediolateral and anteroposterior loading rates [30]. This results in similar resultant loading rates (algebraic sum of the vertical, mediolateral, and anteroposterior) between RFS and FFS habituated to conventional shoes. However, FFS runners habituated to minimal shoes demonstrate significant reductions in vertical, anteroposterior, mediolateral, and thus resultant loading rates [30]. Therefore, if the goal is to reduce overall impact loading, forefoot striking in minimal shoes with adequate time for transition is recommended.

39.5

Effect of minimal shoes on the foot musculoskeletal system

Investigators have assessed the effect of minimal shoes, across a variety of activities, on foot intrinsic muscle strength. Two studies have focused on the use of minimal shoes during training programs. The first study was conducted by Bru¨ggeman in 2005 following the emergence of the Nike Free shoe, a shoe designed to mimic barefoot running [70]. While it had a midsole and therefore was not fully minimal, the shoe had a highly flexible sole and heel counter and provided no arch support. Bru¨ggeman compared the effect of wearing these shoes compared to traditional running shoes on both intrinsic and extrinsic foot muscle strength following a 5-month, warm-up training program. No changes were noted in anatomical cross-sectional areas (ACSA) of the extrinsic foot muscles such as tibialis anterior and posterior, peronei and triceps surae muscles in either group. However, the ACSA increased between 4% and 5% for the intrinsic foot muscles (flexor hallucis, flexor digitorum, abductor hallucis, and quadratus plantae) in the Nike free group. While strength of the dorsiflexors and evertor muscles was unchanged in both groups, metatarsal phalangeal joint flexor, plantarflexor, and invertor strength increased in the intervention group. It was interesting to note that balance and agility were also significantly increased in those that trained in the Nike free shoe. Goldmann et al. [71] assessed toe flexor strength following a high-intensity training program involving jumping, sprinting, and cutting following 15 sessions (5 times/week for 3 weeks). This program involved 5000 6000 push-off maneuvers over the training period. Results between those training in minimal shoes (Nike Free), those training in traditional shoes, and a control group who did not train were compared [71]. The authors reported significant strength gains in both training groups when toe flexion was measured in zero degrees of toe extension. However, only the minimal footwear group demonstrated significant increases when measured in 25 degrees of toe extension. This increased strength at this range was attributed to the greater flexibility of the minimal shoe placing greater demands on the toe flexors. Several other studies have assessed the effect of transitioning to minimal shoes in runners [72 75]. Miller et al. [72] compared the foot characteristics of runners who transitioned to Vibram 5 finger shoes over a 12-week period compared to a control group who remained in traditional running footwear. Using MRI, these authors noted 22% 25% increases in the muscle volumes of the abductor hallicus, flexor hallicus brevis, and abductor digiti minimi in the minimal footwear group. In a similar study incorporating a 6-month transition to Vibram 5 Finger shoes, significant increases in the forefoot intrinsic musculature volume using MRI were noted [75]. Johnson et al. [76] assessed the effect of a 10-week transition program to the Vibram 5 Finger shoes on the ACSA of four muscles using ultrasound imaging. Results suggested a significant increase of 10.6% in the abductor hallucis, but no increases were noted in the other muscles (flexor digitorum brevis, flexor hallicus brevis, and extensor digitorum brevis). Zhang et al. [75] noted a 34% greater arch stiffness and a 9.2% significantly larger abductor hallicus in runners habituated to minimal shoes compared to those in neutral shoes. These results are supported by others who have documented the importance of the abductor hallicus to the integrity of the arch [77,78].

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Habituation to minimal footwear influences the lower leg musculature. Chen et al. [75] noted an increase in the lower leg muscle volume following a 6-month transition to running in minimal footwear. Additionally Histen et al. [79] reported significant changes in the characteristics of the Achilles tendon in runners habituated to minimal footwear (average 4.2 ( 6 1.6) years of minimal footwear running). The Achilles tendons of minimal shoe runners had a 13.5% larger cross-sectional area, were 90.3% stiffer, and could generate 69% greater force than those habituated to conventional shoes. These are all characteristics suggestive of a healthier tendon. The greater capacity of the Achilles is likely related to the greater tendency toward an FFS pattern when in minimal shoes. In fact, a recent study by Wearing et al. [80] reported that habitual FFS runners exhibit significantly greater Achilles tendon stiffness during both walking and running. Therefore, running in minimal shoes can reduce the 52% lifetime incidence of Achilles tendinopathy experienced by male runners [81]. Simply walking in minimal shoes without support increases foot muscle size and strength. In a prospective study, Ridge et al. [82] assessed the baseline size (using ultrasound) of four intrinsic muscles (abductor hallicus, flexor hallicus brevis, flexor digitorum brevis, and quadratus plantae) and three extrinsic muscles (tibialis posterior, tibialis anterior, and flexor digitorum longus). Baseline strength of arch doming and toe flexion was also measured. Subjects were randomized into three groups: a control, a foot exercise group, and a minimal footwear progressive walking group (Inov-8 Bare 210). Following an 8-week intervention, similar, statistically significant, increases in both muscle size and strength were noted in the foot strengthening and minimal footwear group compared to the controls. As compliance with exercise programs can be challenging, these results suggest that minimal shoes can provide a powerful alternative for foot strengthening. A recent cross-sectional study compared the foot strength and arch stiffness of individuals habituated to walking in minimal shoes to those habituated to conventional shoes [83]. As with Ridge et al. [82], these authors noted larger abductor hallucis and abductor digiti minimi muscles (but not flexor digitorum brevis) in the minimally shod group. Additionally, arch height and arch stiffness were greater in this group. Stronger foot muscles also provide greater support to the bones of the foot. The metatarsals, being long and thin, are particularly susceptible to bending moments and increased strain, which can eventually lead to a bone stress injury [84]. However, activation of the plantar foot muscles reduces the bending moments and strain in the second metatarsal [85]. Yet another study indicated that arches deflected less and were more stiff (less bending) in individuals habituated to minimal shoes, likely due to greater activation of the plantar foot muscles [83]. Therefore, the increased strength and activation of the plantar foot musculature may have a positive influence on bone health as well.

39.6

Summary

In summary, barefoot running is our most natural running state. However, barefoot running can be limited by environmental factors such as outside temperatures and surface terrain. Minimal footwear, with its high degree of flexibility and lack of cushioning, allows natural foot motion to occur, while protecting the foot from the elements. Studies have shown that, while not identical to being barefoot, running in minimal footwear results in mechanics that are most similar to barefoot than any other type of running footwear. Removing the support of a traditional shoe has been shown in several studies to result in increases in the intrinsic foot muscle size. These muscles are critical to the integrity of the arch, and when they are deconditioned, can lead to greater strain of the plantar fascia. This strain repeated many times can potentially lead to plantar fasciitis. Being habituated to running in minimal footwear is also associated with stronger, stiffer Achilles tendons compared to conventional running shoes. This may reduce the risk of Achilles tendinopathy. Finally, stronger muscles help to reduce the loading to the bones they support. This potentially could lead to fewer metatarsal stress fractures. However, there are no studies on the effect of minimal footwear use on bone structure or on some of these common running-related injuries. It must be noted that any transition to minimal footwear places new loads on the musculoskeletal system and, like any new activity, should be progressed gradually. Starting children in minimal shoes (both in walking and athletics) removes the need to go through a transition period. It will also lead to a foot and lower leg musculoskeletal system that is stronger and perhaps less prone to overuse injuries.

39.7

Future research

There is still much to be learned about minimal footwear. We need large-scale, longer duration RCTs to fully understand the effect of minimal shoes on mechanics, the musculoskeletal system, and injury risk. For example, we need to better determine what constitutes a true habituated state, as well as the duration of transition to achieve this state. We need to understand the optimal method of transition, and whether it should be augmented with a progressive strengthening program. Future studies should include assessing the effects of starting children in minimal footwear on their

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developing musculoskeletal system. We need to follow cohorts of these children to determine the effect of minimal footwear on their overall foot function and foot health as they become adults. These types of studies will help to provide a stronger rationale for both transitioning adults and starting our children in footwear, which promotes more natural foot function.

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[74] Lindlein K, et al. Improving running economy by transitioning to minimalist footwear: a randomised controlled trial. J Sci Med Sport 2018;21 (12):1298 303. [75] Chen TL, et al. Effects of training in minimalist shoes on the intrinsic and extrinsic foot muscle volume. Clin Biomech (Bristol, Avon) 2016;36:8 13. [76] Johnson AW, et al. The effects of a transition to minimalist shoe running on intrinsic foot muscle size. Int J Sports Med 2016;37(2):154 8. [77] Fiolkowski P, et al. Intrinsic pedal musculature support of the medial longitudinal arch: an electromyography study. J Foot Ankle Surg 2003;42 (6):327 33. [78] Headlee DL, et al. Fatigue of the plantar intrinsic foot muscles increases navicular drop. J Electromyogr Kinesiol 2008;18(3):420 5. [79] Histen K, et al. Achilles tendon properties of minimalist and traditionally shod runners. J Sport Rehabil 2017;26(2):159 64. [80] Wearing SC, et al. Do habitual foot-strike patterns in running influence functional Achilles tendon properties during gait? J Sports Sci 2019;37 (23):2735 43. [81] Maffulli N, Wong J, Almekinders LC. Types and epidemiology of tendinopathy. ClSports Med 2003;22(4):675 92. [82] Ridge ST, et al. Walking in minimalist shoes is effective for strengthening foot muscles. Med Sci Sports Exerc 2019;51(1):104 13. [83] Holowka NB, Wallace IJ, Lieberman DE. Foot strength and stiffness are related to footwear use in a comparison of minimally- vs. conventionally-shod populations. Sci Rep 2018;8(1):3679. [84] Warden SJ, Davis IS, Fredericson M. Management and prevention of bone stress injuries in long-distance runners. J Orthop Sports Phys Ther 2014;44(10):749 65. [85] Sharkey NA, et al. Strain and loading of the second metatarsal during heel-lift. J Bone Jt Surg Am 1995;77(7):1050 7.

Chapter 40

Foot Orthoses Scott Telfer1,2,3 1

Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States, 2Department of Mechanical Engineering,

University of Washington, Seattle, WA, United States, 3RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States

Abstract A foot orthosis is a simple intervention that is often prescribed for a variety of musculoskeletal conditions affecting the foot and other parts of the body. These are used to alter the interface between the foot and the shoe as well as to change some aspect of foot and ankle biomechanics during activities of daily living, such as walking and running. This chapter will describe the design of these devices, their common features, and the materials used in their construction. Kinematic, kinetic, and other biomechanical effects of foot orthoses on the foot and ankle will be reviewed, along with common conditions for which they are prescribed together with their proposed mechanisms of action. Potential areas of future research will be discussed.

40.1

Introduction

Foot orthoses (FOs), also commonly referred to as insoles, orthotics, or shoe inserts, are simple and generally low-risk interventions that are regularly used to treat a variety of musculoskeletal disorders [1,2]. These devices are shaped to fit in the bottom of a shoe and have an upper surface that will conform to varying extents with the plantar surface of the foot. FOs can be made from a variety of materials depending on their intended purpose. Their geometry can be customized to the shape and, to some extent, function of a particular foot [3], or they can be prefabricated “off-the-shelf” devices (Fig. 40.1), intended to approximately fit a range of individuals [4]. The mode of action of most FOs is primarily biomechanical in nature, achieving their effect through kinematic or shock absorbing paradigms [5], although there is some evidence for proprioceptive and neuromuscular effects [6]. This chapter will mainly focus on what are generally described as functional FOs, where the intention is for the device to alter some aspect of the biomechanics of the foot and ankle. Insoles that are primarily intended to provide pressure offloading to reduce the risk of plantar ulcers developing in patients with diabetes are covered in more detail elsewhere in this book, as are ankle-FOs (AFOs).

40.1.1 Design and manufacture of foot orthoses FOs range from fully custom devices that are designed specifically for an individual, to standardized products that may come in a limited number of sizes, with a range of variants in between. The amount of customization required to achieve and maximize a therapeutic effect remains a matter of some debate and likely differs between the conditions for which the FOs are being prescribed [7]. Differences in the level of customization are also reflected in the cost of the devices: off-the-shelf devices can generally be bought in most drug stores for 10 20 USD; whereas fully custom designed and manufactured FOs can cost several hundred USD. Traditionally, the production of custom FOs has been an artisan craft [8]. This process typically requires either a plaster cast or a foam box impression of the patient’s foot to be obtained, which is then used as a negative mold of the foot shape. The negative mold is filled, typically with plaster, to create a positive mold. Once set, the positive mold can be modified by an orthotist through the removal or addition of material. This results in changes to the shape of the final Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00024-X © 2023 Elsevier Inc. All rights reserved.

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FIGURE 40.1 Prefabricated foot orthoses. Figure 1 in Chapman GJ, Halstead J, Redmond AC. Comparability of off the shelf foot orthoses in the redistribution of forces in midfoot osteoarthritis patients. Gait Posture 2016;49:235 240.

device and can create the features required by the prescription (common features are discussed later in this section). The primary FO material, generally a thermoplastic such as polypropylene, is then formed around the corrected positive, trimmed to size, and any further required features added as necessary to produce the final device. This physical casting and molding approach are still used in many locations worldwide; however, in recent years, the implementation of digital design and manufacturing workflows has gained popularity [9,10]. In this case, a typical process requires the plantar foot shape to be digitally captured, either through direct scanning or from a scan of a foam box impression [11], and the scanned geometry transferred to a software program that has tools to allow the FO to be designed around the imported shape. This will usually include options to add similar corrections, modifications, and features that would be created manually using the more traditional process. After a 3D digital model of the finalized design has been created, the software will generate instructions for a computer numerical control milling machine to manufacture the FO from a block of material (direct milling). Alternatively, 3D printing (additive manufacturing) has also recently become an option for producing FOs [10]. Using a fully digital approach for design and manufacture tends to be cleaner and cheaper [12], and furthermore, allows for greater standardization of the design process, allows patient geometry information to be easily transferred, and makes the FO design easy to store and modify in a controlled manner for future prescriptions. In changing the geometry and/or material composition of the FO below the foot, the prescriber is generally attempting to alter the alignment and/or dynamic motion of the bones above, either in the foot or in the lower extremity. There are numerous features of FOs that may be included or adjusted to try to achieve an intended biomechanical effect (Fig. 40.2). Some of the more common of these features are: (1) hindfoot posting [13], where a flat-bottomed post is added beneath the shell of the device and the frontal plane angle of the component is altered medially or laterally to try to change the frontal plane alignment of the calcaneus; (2) heel skives [14], which take the form of a varus or valgus wedge within the heel cup of the FO and are intended to have a similar effect as hindfoot posts; (3) forefoot posting [15], an external modification similar to hindfoot posting but designed to influence the alignment of the forefoot; (4) cutouts in the device intended to accommodate deformities by reducing focused pressure regions or allowing movement of certain joints (for example, this is often used for the first metatarsophalangeal joint [16]); (5) metatarsal domes or bars and forefoot pads are raised regions generally placed proximal to the distal metatarsal heads that can be prescribed to reduce pressure at the forefoot [17]; (6) heel lifts added below the FO that can be used to reduce Achilles tendon load [18] or compensate for leg length discrepancy; and (7) different levels of arch support to compensate for deformity or help to offload over regions of the foot [19]. Custom FOs and their features are primarily designed using measurements and impressions of the static foot, along with clinical assessments that in some cases include semi-quantitative measurements of functional aspects such as the range of motion of a particular joint, or visual observations of walking patterns. Several researchers have proposed

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FIGURE 40.2 Some of the common features found in functional foot orthoses. (A) Hindfoot post; (B) forefoot post; (C) heel skive. (A) Supplied by the author. (B) From Hsu WH, Lewis CL, Monaghan GM, et al. Orthoses posted in both the forefoot and rearfoot reduce moments and angular impulses on lower extremity joints during walking. J Biomech 2014;47:2618 2625. (C) From Bonanno Dr, Zhang CY, Farrugia RC, Bull MG, Raspovic AM, Bird AR, et al. The effect of different depths of medial heel skive on plantar pressures. J Foot Ankle Res. 2012;5:20.

incorporating objective measurements of dynamic foot function in the design of FOs, including data from pressure measurements [20] and multi-segment foot kinematics [21]. This patient-tailored approach facilitates more targeted changes in function; however, there have been no large-scale randomized trials of functional FOs assessing if it can lead to improved clinical outcomes. It is worth noting that in the case of patients with diabetes who are at-risk of developing foot ulcers, there is growing evidence that offloading insoles designed using objective measurements of plantar pressure may be more effective than standard, shape-based insoles at reducing pressures below at-risk areas [22,23] and ultimately preventing ulcer formation [24]. The materials used to construct FOs also play an important role in the biomechanical effects of the device, and this is a potential area for future innovation. Typically, functional FOs use a rigid plastic shell, often polypropylene, as the main component of the device, then a softer top cover layer is added to improve comfort and, in some cases, reduce plantar pressures. Commonly used materials for top covers are open- and closed-cell foams including polyurethane (Poron), polyethyelene (Plastazote), and high-density ethylene-vinyl acetate (EVA). Measurement of in-shoe plantar pressures has allowed researchers to characterize the effects of these materials, and all of them have some pressure reliving properties [25]. These foams, however, do tend to lose their effectiveness over time and with use [26]. Beyond this, several alternative materials have been used in FO designs. Orthoses that incorporate shock absorbing materials, such as gel pockets located below the heel, have been tested in military [27] and sporting populations [28]. The use of textured covers has been suggested to improve proprioception and is being explored as a intervention to reduce the risk of falls in older adults [29]. Novel, 3D printed devices have also been proposed, with this manufacturing approach providing additional design freedom to create devices and features that would not be possible or would be prohibitively expensive with traditional manufacturing processes (Fig. 40.3). For example, this approach allows lattice structures to be created that can be “tuned” to produce different material properties [30], and 3D printing has been used to make multiple versions of the same FOs but with varying overall stiffnesses [31]. Other novel devices that have been suggested include those with embedded sensors to measure plantar forces [32], or temperature [33], as well as those with adjustable components built into the design [10].

40.2

Biomechanical effects of foot orthoses

This section will discuss the biomechanical effects of FOs, along with the methods that have been used to quantify them. A significant proportion of the research into the biomechanical effects of FOs has been carried out in healthy or asymptomatic individuals. The effects of these devices on specific medical conditions are covered in more detail in the

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FIGURE 40.3 Custom foot orthoses produced using a 3D printing process (selective laser sintering of nylon 11 material). Note the geometry of the mesh-like regions which would be difficult to manufacture in this custom device using traditional or subtractive manufacturing methods. The geometry of the mesh at these regions can be varied to change the overall material properties of the FO.

FIGURE 40.4 Adapted footwear used in FO studies to allow markers to be placed directly on the foot and remain in place while the FO condition is changed.

next section. In addition, it should be noted that to determine the effect of FOs in laboratory experiments, footwear is usually standardized, which may partially limit the external validity of findings from these studies.

40.2.1 Kinematic effects of foot orthosis Studying the kinematic effects of FOs on the more distal joints of the limb, particularly the intrinsic foot joints, is challenging. Standard marker-based motion capture protocols, where retroreflective markers are attached to the footwear worn by the participant, is a poor proxy for measuring the actual movement of the foot within the shoe [34]. To address this problem, researchers have modified shoes by cutting windows in them (Fig. 40.4), thus allowing markers to be attached directly to the skin of the foot [35]. This approach can be effective; however, depending on the number and size of the holes, they may affect the structural integrity of the shoe. It also does not overcome issues with skin motion artifacts, where the skin, and therefore the markers mounted on it, move relative to the underlying bone, introducing errors in the measurement [36]. To overcome this, bone pins, where markers are attached to pins surgically inserted into the foot bones, have been used in FO studies to track the “true” bone motion [37,38]. It should be noted that these were small studies (n 5 5) and that the highly invasive nature of the technique severely restricts its applicability to the research realm. Biplane fluoroscopy, where X-ray videos are taken of the foot from two directions and matched to the shape of a 3D bone model to find its position and orientation, has been used to measure kinematic changes caused by FOs [39]. This is a non-invasive method, although it does require the subject to be exposed to ionizing radiation (including a CT scan to produce the bone models), again restricting its applicability outside of research studies. In addition, the complexity and overlapping nature of the foot bones when viewed on planar X-ray can make some bones difficult to track. Terms such as “motion-control” or “anti-pronation” are commonly found in advertising materials for FOs. While these terms tend to be ill-defined and inconsistently used, FOs have several significant effects on the kinematics of the foot and ankle. These results are perhaps strongest for effects on movements in the frontal plane. Medial hindfoot posting has been shown, across several studies, to reduce peak hindfoot eversion in normal [40] and planus foot types [41] during walking. Furthermore, several studies have found dose-response type effects [13,42], showing that by varying the level of posting a systematic change in hindfoot eversion can be achieved, providing strong evidence for the effects of these features (Fig. 40.5). Some have argued that the overall influence of FOs on these kinematics is small; however,

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FIGURE 40.5 Kinematic changes of hindfoot motion with altered levels of hindfoot posting on those with normal foot type (Control) and symptomatic flat foot (Patient). nL: degrees laterally posted; 0 N: neutrally posted; nM: degrees medially posted. Adapted from Telfer S, Abbott M, Steultjens M, Woodburn J. Dose-response effects of customized foot orthoses on lower limb kinematics and kinetics in pronated foot type. J Biomech 2013;46: 1489 95.

it should be noted that the range of motion of the hindfoot is relatively limited outside of the sagittal plane, so relatively small angular changes may be significant in this context. The amount of internal rotation of the tibia in the transverse plane is another kinematic variable that has been suggested to be affected by FOs [43]. This rotation was reduced during running [37]; however, systematic reviews of the literature found no overall evidence to support a change in the peak rotation during walking in normal [40] or flat foot types [41].

40.2.2 Kinetic effects of foot orthosis In terms of kinetics, the majority of studies looking at changes in joint moments induced by FOs have used a single force plate to measure ground reaction forces for each foot. This means that most studies have been limited to studying kinetic effects at the ankle joint complex, and not the intrinsic joints of the foot. In addition, studies looking at joint moments require kinematic information as input data and thus the same limitations as described in the previous section should be noted. Ongoing research using musculoskeletal models may provide an avenue for more detailed changes in intrinsic foot kinetics to be studied [44]. FOs with a medially posted hindfoot reduce peak ankle eversion moment in normal foot types [40] during walking, and similar to kinematics, dose response effects in the frontal plane have been seen when altering the level of hindfoot posting [13]. The effect was less clear for flat foot types [41]. During running, FOs consistently reduce the external ankle eversion moment [45]. Peak loading rates during running have also been found to be reduced with certain types of FOs [46], as well as overall peak ground reaction forces [47]. There is some evidence that FOs can achieve kinetic effects on more proximal joints; these are discussed in the context of knee osteoarthritis later in this chapter.

40.2.3 Effects of foot orthosis on plantar pressure Another form of force measurement, plantar pressures are often measured in studies on the effects of FOs. In-shoe pressure measurement systems tend to be one of the simpler measurement systems to set up, and these generally consist of a thin, pressure sensitive insole containing a matrix of sensors with a density as high as 4 sensels per cm2 that is placed between the FO and the plantar surface of the foot, and give the researcher the ability to map the changes in pressure across the different regions of the plantar foot. While accommodative FOs are designed primarily to reduce the load on certain regions of the foot, functional FOs have also been found to affect plantar pressures. In particular, the addition of an arch support can significantly redistribute load from the forefoot and/or heel to the midfoot region [48]. In addition, the inclusion of features such as metatarsal domes reduces peak plantar pressures under the metatarsal head by 45 60 kPa [17].

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FIGURE 40.6 EMG ensemble average for tibialis posterior and peroneus longus derived from a single gait cycle for all participants. Gray shading represents 95% confidence interval for the shoe-only condition. Figure 5 from Murley GS, Landorf KB, Menz HB. Do foot orthoses change lower limb muscle activity in flat-arched feet toward a pattern observed in normal-arched feet? Clin Biomech 2010;25:728 736.

40.2.4 Effects of foot orthosis on muscle activity patterns The effects of FOs on muscle activation patterns have been studied using electromyography (EMG). Most of the research on the effects of FOs on muscle activity has been on the superficial muscles that actuate the ankle joint complex as these can be measured using electrodes mounted on the skin directly above the muscle (surface EMG). Some work has been performed using fine wire electrodes implanted into the muscle, primarily looking at the tibialis posterior [49]. Generally, there is limited evidence for effects of FOs on muscles acting primarily in the sagittal plane during walking. Murley found evidence that FOs could help produce peroneus longus EMG patterns in people with flat feet that were closer to those usually seen in normal feet (Fig. 40.6) [50]. For running, a systematic review by Reeves et al. found limited evidence that FOs could influence muscle activity in in the sagittal or frontal plane [51], with indications of effects on the tibialis posterior and peroneus longus muscles. In a separate running study, a wavelet-based approach was used to assess the frequency components of the EMG signals for lower limb muscles, finding that FOs tended to increase the intensity of the EMG signal, particularly in the high frequency bands [52]. This led them to suggest that the muscles may have a role in minimizing soft tissue vibrations. FOs may have an effect on the intrinsic muscles of the foot; however, this is challenging to measure. A small study using intramuscular EMG found that FOs reduce the neural drive to the adductor hallucis and first dorsal interosseous muscles; however, further research is required to see if this effect is true for those with clinical conditions [53]. There is also some emerging evidence that intrinsic muscle cross-sectional area can be reduced after wearing FOs for a period of time [54].

40.3

Effects of foot orthosis on clinical conditions

FOs are prescribed for a variety of conditions involving musculoskeletal pain at sites across the body, from the toes to the back. Often, systematic reviews of the efficacy of this type of intervention report that the evidence for their use is weak. This may be in part due to a combination of difficulties in blinding participants to the intervention, the heterogeneity of the devices between different trials, and other factors. This chapter will not delve too deeply into this and will focus mainly on the biomechanical mechanisms that have been proposed and measured in these patient groups.

40.3.1 Rheumatoid arthritis Rheumatoid arthritis often affects the small joints of the hands and feet, leading to pain and deformity at these sites. FOs are regularly prescribed for this condition and have been found in a recent systematic review to be effective at

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reducing plantar pressures in the forefoot region [55], likely leading to reduced pain levels [56]. In cases where the hindfoot is also affected, custom FOs have been shown to help correct the deformity and improve gait kinematics by reducing hindfoot eversion and internal tibial rotation [57].

40.3.2 Symptomatic flat foot FOs reduce hindfoot eversion in people with flat feet [41], likely as a result of providing some correction of the deformity. There is good to moderate evidence that FOs reduce pain, improve function, as well as reducing hindfoot eversion and alter loading patterns in this population [58]. Peak plantar pressures at the heel and forefoot are reduced by both custom and off-the-shelf FOs in this patient population [48].

40.3.3 Heel pain (plantar fasciitis) While FOs are commonly prescribed for heel pain, the mechanism of action is not well defined. There is limited evidence for the devices ability to reducing peak heel pressure in this patient population and therefore offload damaged tissue in the heel [59]. Another suggested mechanism is that the FOs directly reduce the strain on the plantar aponeurosis by supporting the arch [60]. There is evidence that FOs can reduce heel pain in the medium term (7 12 weeks) compared to a sham (placebo) FO; however, it is not clear if this effect is clinically significant, as there was no significant improvement found for longer or shorter timeframes [61].

40.3.4 Osteoarthritis Foot: FOs reduce hindfoot and forefoot pressures while increasing the force under the midfoot in patients with midfoot arthritis [62]. By supporting the medial longitudinal arch, FOs are thought to reduce the bending moment on the midfoot, and this may explain reductions in pain reported in clinical trials [63]. Knee: Laterally wedged insoles have been proposed as a treatment for medial compartment knee osteoarthritis. The proposed mechanism is that by shifting the center of pressure of the ground reaction force laterally during stance, the external knee adduction moment can be reduced [64], thus reducing loading on the medial component of the knee. There is evidence that lateral wedged FOs can achieve this, with systematic reviews finding evidence for reductions in knee adduction angle and knee adduction moments [65]. However, results from clinical trials do not currently support these biomechanical changes translating into clinical benefits in terms of pain reduction [66]. This may be due to the relatively small reduction in the knee adduction moment, of the order of B5% for FOs with 5 degrees of lateral wedging. Larger wedging angles can be used but may become uncomfortable for the patient to wear. There is also some data that suggests that a subgroup of individuals may be “non-responders” to this type of intervention, perhaps due to differences in the effect of the insoles on ankle kinematics [67].

40.3.5 Sports injuries and other conditions There is some evidence showing that FOs are effective at preventing stress fractures and other non-soft tissue injuries [68], presumably driven by a reduction in load across the bone. Recent randomized control trials have provided some evidence that FOs are effective in reducing overuse injuries, including medial tibial stress syndrome, patellofemoral pain, Achilles tendinopathy, and plantar fasciitis/plantar heel pain, in both sports [69] and military training populations [70]. However, the mechanism of action of these effects is often unclear. There is some evidence of an association between back pain and flat foot posture [71]. Foot posture has been shown to affect pelvic positioning [72] and having a flat foot type may place strain on the sacroiliac and lumbosacral joints. In addition, it has been proposed that high arched feet reduce shock absorption and predispose to shock induced pathology of the lower back [73]. However, a systematic review of FOs for back pain found no evidence for their effectiveness, either for treatment or prevention [74].

40.4

Areas of future research

While a significant amount of research has studied the biomechanical and clinical effects of FOs, several questions still need to be answered. In terms of the effects of FOs, it is still often unclear which patients will respond in terms of biomechanical corrections to the FOs. There are also questions about the potential long-term detrimental effects of

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regularly wearing FOs, with some limited evidence suggesting they weaken the intrinsic muscles of the foot [54], and this needs further investigation. Questions relating to how biomechanical effects relate to pain and other symptoms between and across many conditions remain open. The use of musculoskeletal models to non-invasively explore effects on individual structures may help to provide some answers in these cases. In terms of design and manufacture, 3D printing methods present the opportunity to develop insoles with novel features and regionally variable properties by controlling the structure of the material [30]. The use of embedded sensors in “smart” insoles is gaining some traction [75], and these could be combined with other sensors and linked to smart phones to provide feedback to the wearer. Finally, the use of dynamic instead of static measurements to design FOs presents an interesting avenue for customization.

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[22] Telfer S, Woodburn J, Collier A, Cavanagh PR. Virtually optimized insoles for offloading the diabetic foot: a randomized crossover study. J Biomech 2017;60:157 61. [23] Owings TM, Woerner JL, Frampton JD, Cavanagh PR, Botek G. Custom therapeutic insoles based on both foot shape and plantar pressure measurement provide enhanced pressure relief. Diabetes Care 2008;31:839 44. Available from: https://doi.org/10.2337/dc07-2288. [24] Ulbrecht JS, Hurley T, Mauger DT, Cavanagh PR. Prevention of recurrent foot ulcers with plantar pressure-based in-shoe orthoses: the CareFUL prevention multicenter randomized controlled trial. Diabetes Care 2014;37:1982 9. Available from: https://doi.org/10.2337/dc13-2956. [25] Gerrard JM, Bonanno DR, Bonanno DR, Whittaker GA, Whittaker GA, Landorf KB. Effect of different orthotic materials on plantar pressures: a systematic review. J Foot Ankle Res 2020;13. Available from: https://doi.org/10.1186/s13047-020-00401-3. [26] Cronkwright DG, Spink MJ, Landorf KB, Menz HB. Evaluation of the pressure-redistributing properties of prefabricated foot orthoses in older people after at least 12 months of wear. Gait Posture 2011;34:553 7. Available from: https://doi.org/10.1016/j.gaitpost.2011.07.016. [27] House CM, Dixon SJ, Allsopp AJ. User trial and insulation tests to determine whether shock-absorbing insoles are suitable for use by military recruits during training. Mil Med 2004;169:741 6. Available from: https://doi.org/10.7205/MILMED.169.9.741. [28] Nunns MPI, Dixon SJ, Clarke J, Carre´ M. Boot-insole effects on comfort and plantar loading at the heel and fifth metatarsal during running and turning in soccer. J Sports Sci 2016;34:730 7. Available from: https://doi.org/10.1080/02640414.2015.1069378. [29] Hatton AL, Gane EM, Maharaj JN, Burns J, Paton J, Kerr G, et al. Textured shoe insoles to improve balance performance in adults with diabetic peripheral neuropathy: study protocol for a randomised controlled trial. BMJ Open 2019;9. Available from: https://doi.org/10.1136/bmjopen-2018-026240. [30] Ma Z, Lin J, Xu X, Ma Z, Tang L, Sun C, et al. Design and 3D printing of adjustable modulus porous structures for customized diabetic foot insoles. Int J Light Mater Manuf 2019;2:57 63. [31] Desmyttere G, Leteneur S, Hajizadeh M, Bleau J, Begon M. Effect of 3D printed foot orthoses stiffness and design on foot kinematics and plantar pressures in healthy people. Gait Posture 2020;81:247 53. [32] Sto¨ggl T, Martiner A. Validation of Moticon’s OpenGo sensor insoles during gait, jumps, balance and cross-country skiing specific imitation movements. J Sports Sci 2017;35:196 206. Available from: https://doi.org/10.1080/02640414.2016.1161205. [33] Telfer S, Munguia J, Pallari J, Dalgarno K, Steultjens M, Woodburn J. Personalized foot orthoses with embedded temperature sensing: proof of concept and relationship with activity. Med Eng Phys 2014;36:9 15. Available from: https://doi.org/10.1016/j.medengphy.2013.08.002. [34] Alcantara RS, Trudeau MB, Rohr ES. Calcaneus range of motion underestimated by markers on running shoe heel. Gait Posture 2018;63:68 72. Available from: https://doi.org/10.1016/j.gaitpost.2018.04.035. [35] Bishop C, Arnold JB, Fraysse F, Thewlis D. A method to investigate the effect of shoe-hole size on surface marker movement when describing in-shoe joint kinematics using a multi-segment foot model. Gait Posture 2015;41:295 9. Available from: https://doi.org/10.1016/j.gaitpost. 2014.09.002. [36] Shultz R, Kedgley AE, Jenkyn TR. Quantifying skin motion artifact error of the hindfoot and forefoot marker clusters with the optical tracking of a multi-segment foot model using single-plane fluoroscopy. Gait Posture 2011;34:44 8. Available from: https://doi.org/10.1016/j.gaitpost. 2011.03.008. [37] Stacoff A, Reinschmidt C, Nigg BM, van den Bogert AJ, Lundberg A, Denoth J, et al. Effects of foot orthoses on skeletal motion during running. Clin Biomech (Bristol, Avon) 2000;15:54 64. Available from: http://www.ncbi.nlm.nih.gov/pubmed/10590345 accessed 11.10.10]. [38] Liu A, Nester CJ, Jones RK, Lundgren P, Lundberg A, Arndt A, et al. Effect of an antipronation foot orthosis on ankle and subtalar kinematics. Med Sci Sports Exerc 2012;44:2384 91. Available from: https://doi.org/10.1249/MSS.0b013e318265df1d. [39] Balsdon M, Dombroski C, Bushey K, Jenkyn TR. Hard, soft and off-the-shelf foot orthoses and their effect on the angle of the medial longitudinal arch: a biplane fluoroscopy study. Prosthet Orthot Int 2019;43:331 8. Available from: https://doi.org/10.1177/0309364619825607. [40] Hajizadeh M, Desmyttere G, Carmona JP, Bleau J, Begon M. Can foot orthoses impose different gait features based on geometrical design in healthy subjects? A systematic review and meta-analysis. Foot 2020;42. Available from: https://doi.org/10.1016/j.foot.2019.10.001. [41] Desmyttere G, Hajizadeh M, Bleau J, Begon M. Effect of foot orthosis design on lower limb joint kinematics and kinetics during walking in flexible pes planovalgus: a systematic review and meta-analysis. Clin Biomech 2018;59:117 29. Available from: https://doi.org/10.1016/j. clinbiomech.2018.09.018. [42] Telfer S, Abbott M, Steultjens M, Rafferty D, Woodburn J. Dose-response effects of customised foot orthoses on lower limb muscle activity and plantar pressures in pronated foot type. Gait Posture 2013;38:443 9. Available from: https://doi.org/10.1016/j.gaitpost.2013.01.012. [43] McPoil T, Cornwall M. The effect of foot orthoses on transverse tibial rotation during walking. J Am Pod Med Assoc 2000;90:2 11. [44] Oosterwaal M, Telfer S, Torholm S, Carbes S, van Rhijn LW, Macduff R, et al. Generation of subject-specific, dynamic, multisegment ankle and foot models to improve orthotic design: a feasibility study. BMC Musculoskelet Disord 2011;12:256. Available from: https://doi.org/ 10.1186/1471-2474-12-256. [45] McMillan A, Payne C. Effect of foot orthoses on lower extremity kinetics during running: a systematic literature review. J Foot Ankle Res 2008;1:13. Available from: https://doi.org/10.1186/1757-1146-1-13. [46] Mu¨ndermann A, Nigg BM, Humble RN, Stefanyshyn DJ. Foot orthotics affect lower extremity kinematics and kinetics during running. Clin Biomech 2003;18:254 62. Available from: https://doi.org/10.1016/S0268-0033(02)00186-9. [47] Eslami M, Begon M, Hinse S, Sadeghi H, Popov P, Allard P. Effect of foot orthoses on magnitude and timing of rearfoot and tibial motions, ground reaction force and knee moment during running. J Sci Med Sport 2009;12:679 84. Available from: https://doi.org/10.1016/j.jsams. 2008.05.001. [48] Khodaei B, Saeedi H, Jalali M, Farzadi M, Norouzi E. Comparison of plantar pressure distribution in CAD CAM and prefabricated foot orthoses in patients with flexible flatfeet. Foot 2017;33:76 80. Available from: https://doi.org/10.1016/j.foot.2017.07.002.

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[49] Barn R, Turner DE, Rafferty D, Sturrock RD, Woodburn J. Tibialis posterior tenosynovitis and associated pes plano valgus in rheumatoid arthritis: electromyography, multisegment foot kinematics, and ultrasound features. Arthritis Care Res (Hoboken) 2013;65:495 502. Available from: https://doi.org/10.1002/acr.21859. [50] Murley GS, Landorf KB, Menz HB. Do foot orthoses change lower limb muscle activity in flat-arched feet toward a pattern observed in normal-arched feet? Clin Biomech 2010;25:728 36. [51] Reeves J, Jones R, Liu A, Bent L, Plater E, Nester C. A systematic review of the effect of footwear, foot orthoses and taping on lower limb muscle activity during walking and running. Prosthet Orthot Int 2019;43:576 96. Available from: https://doi.org/10.1177/0309364619870666. [52] Mu¨ndermann A, Wakeling JM, Nigg BM, Humble RN, Stefanyshyn DJ. Foot orthoses affect frequency components of muscle activity in the lower extremity. Gait Posture 2006;23:295 302. Available from: https://doi.org/10.1016/j.gaitpost.2005.03.004. [53] Collins N, Salomoni S, Elgueta Cancino E, Tucker K, Hodges P. Foot orthoses induce immediate changes in intrinsic foot muscle EMG activity during walking. J Sci Med Sport 2019;22:S17 18. [54] Protopapas K, Perry SD. The effect of a 12-week custom foot orthotic intervention on muscle size and muscle activity of the intrinsic foot muscle of young adults during gait termination. Clin Biomech 2020;78. Available from: https://doi.org/10.1016/j.clinbiomech.2020.105063. [55] Tenten-Diepenmaat M, Dekker J, Heymans MW, Roorda LD, Vliet Vlieland TPM, Van Der Leeden M. Systematic review on the comparative effectiveness of foot orthoses in patients with rheumatoid arthritis. J Foot Ankle Res 2019;12. Available from: https://doi.org/10.1186/s13047-019-0338-x. [56] Frecklington M, Dalbeth N, McNair P, Gow P, Williams A, Carroll M, et al. Footwear interventions for foot pain, function, impairment and disability for people with foot and ankle arthritis: a literature review. SemArthritis Rheumatism 2018;47:814 24. Available from: https://doi.org/ 10.1016/j.semarthrit.2017.10.017. [57] Woodburn J, Helliwell PS, Barker S. Changes in 3D joint kinematics support the continuous use of orthoses in the management of painful rearfoot deformity in rheumatoid arthritis. J Rheumatol 2003;30:2356 64. Available from: http://www.ncbi.nlm.nih.gov/pubmed/14677177 Accessed 13 Nov 2012. [58] Banwell HA, Mackintosh S, Thewlis D. Foot orthoses for adults with flexible pes planus: a systematic review. J Foot Ankle Res 2014;7:23. Available from: https://doi.org/10.1186/1757-1146-7-23. [59] Bonanno DR, Landorf KB, Menz HB. Pressure-relieving properties of various shoe inserts in older people with plantar heel pain. Gait Posture 2011;33:385 9. Available from: https://doi.org/10.1016/j.gaitpost.2010.12.009. [60] Kogler GF, Solomonidis SE, Paul JP. Biomechanics of longitudinal arch support mechanisms in foot orthoses and their effect on plantar aponeurosis strain. Clin Biomech 1996;11:243 52. Available from: https://doi.org/10.1016/0268-0033(96)00019-8. [61] Whittaker GA, Munteanu SE, Menz HB, Tan JM, Rabusin CL, Landorf KB. Foot orthoses for plantar heel pain: a systematic review and metaanalysis. Br J Sports Med 2018;52:322 8. Available from: https://doi.org/10.1136/bjsports-2016-097355. [62] Chapman GJ, Halstead J, Redmond AC. Comparability of off the shelf foot orthoses in the redistribution of forces in midfoot osteoarthritis patients. Gait Posture 2016;49:235 40. [63] Halstead J, Chapman GJ, Gray JC, Grainger AJ, Brown S, Wilkins RA, et al. Foot orthoses in the treatment of symptomatic midfoot osteoarthritis using clinical and biomechanical outcomes: a randomised feasibility study. Clin Rheumatol 2016;35:987 96. Available from: https://doi. org/10.1007/s10067-015-2946-6. [64] Yasuda K, Sasaki T. The mechanics of treatment of the osteoarthritic knee with a wedged insole. Clin Orthop Relat Res 1987;215:162 72. Available from: http://www.ncbi.nlm.nih.gov/pubmed/3802634 Accessed 30 Nov 2017. [65] Shaw KE, Charlton JM, Perry CKL, de Vries CM, Redekopp MJ, White JA, et al. The effects of shoe-worn insoles on gait biomechanics in people with knee osteoarthritis: a systematic review and meta-analysis. Br J Sports Med 2017. Available from: https://doi.org/10.1136/bjsports2016-097108. [66] Parkes MJ, Maricar N, Lunt M, LaValley MP, Jones RK, Segal NA, et al. Lateral wedge insoles as a conservative treatment for pain in patients with medial knee osteoarthritis: a meta-analysis. JAMA 2013;310:722 30. Available from: https://doi.org/10.1001/jama.2013.243229. [67] Chapman GJ, Parkes MJ, Forsythe L, Felson DT, Jones RK. Ankle motion influences the external knee adduction moment and may predict who will respond to lateral wedge insoles?: An ancillary analysis from the SILK trial. Osteoarthr Cartil 2015;23:1316 22. Available from: https:// doi.org/10.1016/j.joca.2015.02.164. [68] Bonanno DR, Landorf KB, Munteanu SE, Murley GS, Menz HB. Effectiveness of foot orthoses and shock-absorbing insoles for the prevention of injury: a systematic review and meta-analysis. Br J Sports Med 2017;51:86 96. Available from: https://doi.org/10.1136/bjsports-2016-096671. [69] Hirschmu¨ller A, Baur H, Mu¨ller S, Helwig P, Dickhuth HH, Mayer F. Clinical effectiveness of customised sport shoe orthoses for overuse injuries in runners: a randomised controlled study. Br J Sports Med 2011;45:959 65. Available from: https://doi.org/10.1136/bjsm.2008.055830. [70] Bonanno DR, Murley GS, Munteanu SE, Landorf KB, Menz HB. Effectiveness of foot orthoses for the prevention of lower limb overuse injuries in naval recruits: a randomised controlled trial. Br J Sports Med 2018;52:298 302. Available from: https://doi.org/10.1136/bjsports-2017-098273. [71] Menz HB, Dufour AB, Riskowski JL, Hillstrom HJ, Hannan MT. Foot posture, foot function and low back pain: the Framingham Foot Study. Rheumatol (Oxf) 2013;52:2275 82. Available from: https://doi.org/10.1093/rheumatology/ket298. [72] Betsch M, Schneppendahl J, Dor L, Jungbluth P, Grassmann JP, Windolf J, et al. Influence of foot positions on the spine and pelvis. Arthritis Care Res 2011;63:1758 65. Available from: https://doi.org/10.1002/acr.20601. [73] Bird AR, Payne CB. Foot function and low back pain. Foot 1999;9:175 80. [74] Chuter V, Spink M, Searle A, Ho A. The effectiveness of shoe insoles for the prevention and treatment of low back pain: a systematic review and meta-analysis of randomised controlled trials. BMC Musculoskelet Disord 2014;15:140. Available from: https://doi.org/10.1186/1471-2474-15-140. [75] Hegde N, Bries M, Swibas T, Melanson E, Sazonov E. Automatic recognition of activities of daily living utilizing insole-based and wrist-worn wearable sensors. IEEE J Biomed Heal Inform 2018;22:979 88.

Chapter 41

Ankle-Foot Orthoses and Rocker Bottom Shoes Elizabeth Russell Esposito1,2 1

DoD-VA Extremity Trauma and Amputation Center of Excellence, Joint Base San Antonio Fort Sam Houston, San Antonio, TX, United States,

2

Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States

Abstract Ankle-foot orthoses and rocker bottom shoes are commonly prescribed for musculoskeletal conditions that lead to gait impairments. They can help with several factors related to biomechanical function, including, but not limited to, compensating for muscle weakness, correcting deformity, and promoting forward progression. This can lead to improvements in function and quality of life for the wearer. This chapter will describe the biomechanical effects of these devices, their use in musculoskeletal conditions, and different designs available. Potential areas of future research will be discussed.

41.1

Introduction

Patients who experience deformities or weakened capabilities at the ankle joint due to injuries or pathologies, such as strokes, hemiplegia, spinal cord injuries, ankle arthritis, etc. often exhibit gait deficiencies. Depending on the affected musculature, the etiology of these deficiencies can be classified as either weakened dorsiflexors or weakened plantarflexors. Weakened dorsiflexors result in a steppage gait pattern characterized by foot slap during loading response where the foot falls uncontrolled to the ground, producing the distinct “foot slap” sound [1]. Subsequently, foot drop/toe drag may occur throughout the swing phase, preventing proper advancement of the limb and increasing the risk of tripping [1]. To compensate for the reduced ankle dorsiflexion to clear the foot during swing, a hip strategy may be used that incorporates pelvic hiking or circumduction [2]. Conversely, weakened plantarflexors can result in a decreased internal ankle plantarflexor torque at push off in terminal stance and can affect stability during midstance. Orthotic and footwear interventions can be effective for several conditions to (1) establish a safer and more efficient gait, (2) correct poor skeletal alignment, (3) provide a stable base of support, (4) promote forward progression, (5) strengthen weak muscles and control muscular imbalances, and (6) prevent painful ranges of motion.

41.2

Ankle-foot orthoses

A conventional approach for treating weakness of the dorsiflexors and/or plantarflexors is a mechanical brace called an ankle-foot orthosis (AFO). AFOs can be used to modify structural or functional characteristics of the ankle-foot complex in an affected limb and can restore lost ankle function by avoiding painful joint positions, provide support to limbs with decreased strength, and offer energy storage and return properties [3 5]. AFOs are commonly prescribed to compensate for limb impairments, restore mobility and limit adverse effects on gait biomechanics. AFOs are included in rehabilitation plans to improve function following neuromuscular disorders that create residual physical deficits including impaired walking, mobility, balance, and others. They assist with joint deformities, optimize anatomic alignment, provide support, and assist or limit joint motion. AFOs are also recommended to minimize gait dysfunction caused by abnormal muscle function. Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00016-0 © 2023 Elsevier Inc. All rights reserved.

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AFOs function by influencing the joint moments at the ankle and/or foot. The AFO produces an external moment to influence the internal moments generated by the active (e.g., muscle) and passive (e.g., ligaments) structures within the body. There are several ways by which this may occur. AFOs may control rotational and/or translational motion about a joint. They may also control axial forces across a joint or alter the line of action of the ground reaction force [6].

41.2.1 Controlling rotational motion AFOs are primarily prescribed to control rotational motion about a joint. The majority of motion at the ankle is in the sagittal plane during walking but, depending on the type of AFO, it may control motion in multiple planes. For example, a very rigid AFO will often restrict motion about all three planes. The external moment applied by the AFO can also be used to position the joint or a segment into a more optimal alignment. In this case, a patient with drop foot may use an AFO to position the ankle in a relatively fixed alignment close to 90-degrees. Restricting motion about a joint nearly always has effects on the neighboring joints, most predominately the knee joint, but may be tolerated by the patient and provider to promote mobility and function, and/or reduce painful range of motion.

41.2.2 Controlling translational motion Although some translational motion is inherent in a healthy musculoskeletal system, large amounts of translational motion can be considered as deleterious. Instability of a joint (often via loss of ligamentous integrity through degenerative or traumatic means) may lead to adverse translational motions at that joint. An AFO may be used to mitigate translational motion at the ankle using straps, for example, to brace the foot segment with the leg segment, and vice versa.

41.2.3 Controlling axial forces Some AFOs are designed to be “offloading” braces to reduce axial loading that is often painful and debilitating. These AFOs redistribute a portion of the load to structures proximal to the painful joint. A patellar tendon bearing AFO is an example of an offloading design in which the tibial cuff bears a portion of the axial force that would otherwise be required of the ankle.

41.2.4 Altering the line of action of the ground reaction force The line of action of the ground reaction forces during the stance phase of gait creates a moment about every joint in the lower extremity. An AFO can be used to shift how the ground reaction force vector passes an affected joint. This is the primary purpose of an in-shoe foot orthosis. However, the same principles can be incorporated into the design of an AFO by contouring the shape of the footplate to affect this line of action and alter how the foot rolls over the ground. For example, a heel rocker may effectively reduce the internal knee extensor moment when wearing a solid AFO.

41.3

Rocker bottom shoes

Rocker bottom shoes are another means of influencing how the foot rolls over the ground. Rocker bottom shoes are an inexpensive and noninvasive alternative treatment that may be prescribed for patients needing to offload portions of the foot or for those with dorsiflexor or plantarflexor weakness. The curved rocker bottom of the shoe leads to better rolling of the body over the ground and limits the need for foot and ankle movement for forward progression [1]. The rocking motion between heel strike and toe-off redistributes force patterns under the foot and may offload areas of undesirable pressure. They are the most commonly prescribed footwear modification to address a variety of clinical pathologies [7]. The curved rocker sole may improve mechanical efficiency by reducing the positive work required at the ankle at pushoff and also limiting the negative work performed during other phases of the gait cycle (e.g., preload and collision). These reductions result in less net positive work on the body center of mass during a stride [8]. Certain designs of rocker bottom shoes can also favorably reduce forefoot plantar pressures during walking, which may potentially improve healing time after an injury or surgery [9]. They can reduce the need for contributions from the affected musculature and can reduce sagittal plane motion [10,11] that may otherwise be painful for some patient populations and limit function.

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41.4

649

Roll-over shape

Biomechanically, rocker bottom shoes facilitate the advancement of the limb and whole body center of mass over the base of support. With regards to kinematics and kinetics, most studies have suggested that rocker bottom shoes have the greatest effects at the ankle. As the foot and lower limb progress from heel strike through toe-off, the body’s center of mass moves over three sequential anatomical rockers: the heel rocker, the ankle rocker, and the forefoot rocker [1]. The heel incorporates the rounded calcaneal tuberosity that, upon initial loading response, assists in rolling the foot towards the ground. Next, the ankle then rolls the distal tibial-fibular complex across the rounded surface of the talar dome to progress the whole body center of mass over the foot. Finally, the forefoot rocker occurs as the foot rolls over the metatarsophalangeal joints and the toes dorsiflex in terminal stance/preswing. The integrated effect of the ankle-foot rockers is represented by the roll-over shape [1] (Fig. 41.1). Roll-over shape is the geometry to which the ankle-foot conforms during the stance phase of gait [13]. Mathematically, roll-over shape is represented by tracking the trajectory of the center of pressure under the foot in the lower leg segment coordinate system and is viewed by plotting the vertical center of pressure displacement relative to the horizontal displacement. An effective roll-over shape can facilitate forward progression. For decades, researchers have modeled the foot and ankle using rockers (Fig. 41.2). Simulations of rocker models of the foot have shown the sensitivity of the model to the effective shape of the foot [14,15]. Rockers have also been used to describe the function of the ankle-foot complex in walking toys with circular feet of different radii, with greater rocker foot radii improving stability [16]. Combinations of computer and physical models have estimated the optimal rocker radius for human walking (0.3 x leg length) [17]. The importance of the effective rocker shapes of prosthetic feet have also been evaluated [18]. In fact, one of the original purposes of evaluating roll-over shape was to guide the design and alignment of various prosthetic device components [13,19,20] and characterize different prosthetic feet [21]. The same concepts also apply to AFOs (Fig. 41.1) and different footwear conditions (e.g., rocker bottom shoes). In addition, improved ankle-foot kinematics (i.e., through the use of assistive devices such as AFOs or rocker bottom shoes) may also result in a more biomimetic roll-over shape. A “normal” roll-over shape is likely a desirable outcome for patients with pathologies or complications affecting the ankle-foot complex. For example, Fatone et al. [22] found that AFOs improved roll-over shape in stroke patients to the extent that components of roll-over shape were not significantly different from unaffected controls. In a separate study on patients with a variety of lower limb impairments (e.g., stroke, nerve injury, etc.), the authors found that, although clinically prescribed AFOs improved (increased) the radius of curvature compared to walking without an AFO, deficiencies still existed relative to able-bodied individuals [23]. Rocker bottom shoes are specifically designed to facilitate smooth progression of the body over the base of support, similar to a wheel rolling over its curved surface. It is possible that they may also improve stability as mechanical models with rockers can tolerate greater perturbations without falling over than models without rockers [24].

41.5

Patient populations

AFOs and rocker bottom shoes may be prescribed to treat a wide variety of pathologies. They may be used on a temporary basis, or permanently. They may be used throughout all activities of daily living, or only to assist function and

Roll-over shape of the ankle-foot system

Shoe

AFO

Roll-over shape of the orthocankle-foot system

Shoe

FIGURE 41.1 Roll-over shape of the biologic ankle-foot system (left) and with an ankle-foot orthosis (right) [12]. Reproduced from Hansen AH. The quest to improve rocker effects in O&P, Low Extrem Rev 2010.

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FIGURE 41.2 The center of mass passes over the base of support. The integrated effects of the heel, ankle, and forefoot rockers in addition to the rockers provided externally by the geometry of an AFO or shoe.

mobility during certain tasks. Although a comprehensive assessment of walking biomechanics in all patient populations who may benefit from AFOs and rocker bottom shoes is beyond the scope of this chapter, a brief overview of a few selected pathologies is included below. These conditions were selected due to the broad applicability of literature on the topic as well as the clinical acceptability of AFOs and/or rocker bottom shoes as treatment for biomechanical limitations.

41.5.1 Stroke Muscle weakness, spasticity, compromised sensorimotor control, and/or the loss of cognitive functions can all impair walking ability after a stroke [25]. Regaining the ability to walk is a major rehabilitation goal [26]. AFOs are commonly prescribed after stroke to facilitate some of the aspects of normative gait that have been lost as a result of pathology. AFOs can correct joint alignments, prevent the foot drop that can result from a stroke, increase walking speed, and reduce energy expenditure [27 32]. Many studies have been conducted on the efficacy of AFOs on hemiplegic gait and most indicate that AFOs improve function and mobility. A review of 43 studies indicated that AFOs of many different types (e.g., plastic, rigid, articulated, or posterior leaf design, to name a few) resulted in faster walking speeds, greater step and stride length, and greater weight bearing through the affected limb during standing over the short term [33]. Despite some reported benefits, it remains unknown if AFOs consistently result in better balance, less postural sway, and greater overall mobility, and if these can be achieved long-term [33,34]. Some researchers report that the use of an AFO to treat the gait

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impairments that result from stroke may prolong dependence on the assistive device and lead to a weakening of the surrounding musculature, particularly the dorsiflexors [35]. Foot drop, or equino-varus foot, is perhaps one of the most common impairments that affects walking ability. This condition shifts the point of foot contact with the ground from the heel to the lateral plantar surface of the foot and to the mid and forefoot regions, resulting in a potential loss of balance and reduced stride length. A conventional or offthe-shelf AFO (described below) may be used to maintain the ankle in a more dorsiflexed position throughout the gait cycle, supporting it through swing to prevent tripping. This design of AFO inherently limits the peak internal ankle plantarflexor moment at push-off by preventing substantial amounts of ankle plantarflexion to occur but the tradeoff is generally desirable for aspects of function, mobility, and safety during walking. Rocker bottom shoes may be used in combination with AFOs post-stroke. The three rockers described previously— the heel, ankle and forefoot rockers—can be seriously damaged after stroke [1]. AFOs can improve the effect of the heel and ankle rockers during walking but do not appear to affect the forefoot rocker in hemiplegic gait [36,37]. Farmani et al. [38] compared the use of AFOs with standard shoes to AFOs with rocker bottom shoes in hemiplegic patients. The authors found that AFOs used with rocker bottom shoes, as opposed to standard shoes, improved performance on function and mobility tasks, increased walking speed, and reduced the metabolic cost of walking in patients who had suffered from a stroke.

41.5.2 Cerebral palsy Roughly two-thirds of children with cerebral palsy have the ability to walk; however, due to the effects of the condition on the musculoskeletal system, gait can be severely affected. Specifically, over 80% of children with cerebral palsy experience spasticity and imbalances in muscle tone. The contractures that result from associated with cerebral palsy can sometimes be addressed, at least in part, with AFOs. Although several studies have reported improvements in various gait deficiencies with AFOs, a review of the literature up to the year 2000 indicated that AFOs were not strongly supported for children with cerebral palsy [39], potentially due to high variability in designs and low levels of scientific evidence relating to quality of methods. Updated reviews have been similarly limited but have found generally positive effects of AFO use on ankle range of motion, gait kinetics and kinematics, and functional activities in children with cerebral palsy [40]. The magnitude of the gait deficiency largely depends on the type and severity of cerebral palsy. Diplegia is the most common presentation and is often characterized by a slow speed and gait deviations due to contracture on the spastic plantarflexors [41]. Excessive ankle plantarflexion from equinus contracture leads to a premature peak in the internal ankle plantarflexion moment at push-off, excessive knee flexion, and abnormal knee extensor moments during stance [42,43]. To address the gait deviations, over 50,000 AFOs are prescribed each year to provide external support. These AFOs may be solid or hinged. While both designs often improve walking speed and stride length, a hinged AFO allows for greater power generation and more normal dorsiflexion at terminal stance [44]. A solid AFO limits the normal forward advancement of the tibia over the foot and decreases ankle dorsiflexion [44,45]. Both designs are effective in this patient population for reducing excessive ankle plantarflexion without altering knee kinematics during stance. It is important to note that AFOs are rarely used in isolation with cerebral palsy and other therapeutic modalities (e.g., botulinum A toxin, orthopedic intervention, physical therapy intervention) are often used in combination. Reducing energy expenditure is also a key consideration for individuals with cerebral palsy as fatigue can limit participation in activities of daily living. In some cases, the application of an AFO is effective at reducing energy expenditure [46] but there is wide variability across studies. Patients should be independently evaluated as to the efficacy of an AFO on energy expenditure. Mass added distally to the limb, as in the case of an AFO, typically has detrimental effects on energy expenditure in healthy individuals [47]. Cases where energy expenditure decreases or even stays the same with the addition of an AFO may be regarded as a positive outcome provided there are other improvements in gait, mobility, or function.

41.5.3 Ankle arthritis The pain, disability, and reduced ankle range of motion that often result from ankle arthritis can also be addressed with assistive devices. Patients with this ankle arthritis often experience reduced range of ankle motion and are limited in the distance they can walk [48]. Rocker bottom shoes may reduce pain, disability, and activity limitations in patients with rheumatoid arthritis [49]. The benefit of rocker bottom shoes in this patient population is that the shape of the rocker sole allows the body’s center of mass to roll over the base of support without requiring large amounts of sagittal plane motion [10,11,50]. They are routinely prescribed as a noninvasive conservative treatment for ankle arthritis and also

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after ankle fusion [51]. Despite their relatively common prescription for this patient population, the literature is somewhat sparse on the efficacy of rocker bottom shoes as an intervention. Results are somewhat mixed depending on the variables investigated. Van Engelen et al. [52] performed one of the few studies investigating rocker bottom shoes in patients with ankle arthrodesis. The authors compared rocker bottom shoes (Masai Barefoot Technology) to standard walking shoes and found that the rocker bottom shoes did not improve metabolic cost, step-to-step transition work, or roll-over shape. In patients with rheumatoid arthritis, rocker bottom shoes improved Foot Function Index scores [49].

41.5.4 Limb salvage Lower limb salvage, or reconstruction, may be a feasible option after high-energy, open fractures of the distal tibia, ankle, or foot. Gait deviations are common in individuals with lower limb salvage and include slower walking speeds, foot drop, uneven step length, knee hyperextension, hip hiking, Trendelenburg gait, and leg circumduction [53]. AFOs may be prescribed to restore lost function, stabilize the injured joint, and/or prevent painful ranges of motion [54 56]. AFOs may be prescribed during the rehabilitation process to offload the joint during recovery. A military study investigated the use of a dynamic AFO during rehabilitation and found that its use, combined with a rehabilitation initiative, improved physical performance across a variety of outcome measures [57]. Performance improved in patient groups with both ankle fusions and ankle fusions combined with subtalar fusion.

41.6

Design and prescription of ankle-foot orthosis

AFOs vary in their size, shape, material properties, and functional characteristics depending on the clinical application. Well established information on the design and prescription of AFOs for a given patient is still limited. Although the use of AFOs for addressing movement and functional limitations seems to be embraced in recent years [58,59], they are still not consistently prescribed by the provider or used by the patient [60].

41.6.1 Conventional versus advanced AFOs can generally be classified into conventional/off-the-shelf/thermoplastic designs (Figs. 41.3 and 41.4) that may be more appropriate for lower impact activities or “advanced”/dynamic-response/energy-storing-and-returning designs (Fig. 41.5) that may be more appropriate for higher impact activities. Designs may vary in material construction and included componentry. The effects of these devices may be within any plane of motion. Conventional and off-the-shelf AFOs include designs such as conventional double-upright designs, posterior leaf springs, gauntlet, and off-the-shelf carbon fiber designs. Advanced AFOs include designs that are widely believed to return stored energy in late stance. Here we describe them as the category of custom carbon fiber AFOs that are “passive-dynamic,” “dynamic response,” or “energy-storing-

FIGURE 41.3 Hinged and solid thermoplastic ankle foot orthoses (AFOs). (A) and (B) represent articulated AFOs and (C) represents a posterior leaf spring.

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FIGURE 41.4 Off-the-shelf, carbon fiber AFOs: (A) Blue Rocker (Allard), (B) SpryStep Plus (Thuasne), (C) Matrix (Trulife), and (D) WalkOn Reaction (Ottobock).

FIGURE 41.5 Examples of “advanced” custom, carbon fiber AFOs considered dynamic-response or energy-storing-and-returning. Common components include a rigid footplate, a tibial cuff, and a connecting posterior strut. (A) Intrepid Dynamic Exoskeletal Orthosis (IDEO) and (B) Reaktiv (Fabtech).

and-returning (ESR)” designs. These AFOs rely on design features such as material properties, device shape, springs, or fluid pressure dynamics to regulate the storage and return of mechanical energy. Advanced AFOs are generally used in more active patients who participate in high intensity activities. These designs are laminated and may include posterior struts. Due to the complexity of fabrication, individualized alignments across multiple planes, and the need for a very precise fit, prefabrication is not currently utilized, as it often is with conventional AFO designs. ESR AFOs usually incorporate modular components such as struts, brackets, and/or closure systems, but custom fabrication is still required to produce the bulk of most current designs (Fig. 41.5).

41.6.2 Articulated versus nonarticulated AFOs may also fall into categories of articulated or nonarticulated (Fig. 41.3). Fixing the ankle with a nonarticulated AFO requires biomechanical adaptations to maintain walking speed and forward propulsion. For example, a nonarticulated AFO will prolong the contact phase of gait. Because as the ankle cannot plantarflex to lower the foot to the ground, the knee experiences a greater external knee flexor moment to rotate the tibia forward and achieve foot flat on

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the ground. Heel rockers within the shoe or a cushioned wedge under the heel may mitigate the increase in moment to promote better roll-over characteristics. A systematic review on patients with gait impairments from stroke indicate that AFOs improve walking speed regardless of whether articulated AFOs or rigid nonarticulated AFOs are used [61].

41.7

Design and prescription of rocker bottom shoes

Rocker sole design can influence gait and plantar pressure. The design of the rocker sole may include the geometry of the rocker bottom or the material composition. The geometry of the rocker bottom may include features such as the radius of curvature, the length of the rocker, the location of the rocker along the shoe, rocker angle with respect to the ground and with respect to the long axis of the shoe, the shoe height, and the rocker axis position, to name a few [62]. The material of the sole can range from soft to rigid depending on the type of design [7]. The deformation experienced with a soft rocker sole may not appropriately limit sagittal plane ankle joint motion (4,7,19) and a more rigid rocker sole may be desired if the goal is to limit painful ranges of ankle joint motion or to stabilize the ankle joint to promote healing. Traditionally, rocker bottom shoes have a rigid sole with a single rocker at the forefoot. More recently, rocker bottom shoes have incorporated a wider variety of designs including toe-only rockers, negative heel rockers, and double rockers (Fig. 41.6). The design is driven by the specific pathology or weakness they are addressing. For example, a negative heel may be inappropriate for patients with poor ankle dorsiflexion [63]. Unstable rocker bottom shoes are a relatively new use of a rocker sole (e.g., Masai Barefoot Technology, MBT). These shoes provide an unstable base via a rounded rocker in the anterior-posterior direction and are reported to

FIGURE 41.6 Types of rocker bottom shoes [63]. Recreated from Brown D, Wertsch JJ, Harris GF, Klein J, Janisse D, Effect of rocker soles on plantar pressures. Arch Phys Med Rehabil 2004;85(1):81 6.

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promote instability for training purposes (e.g., muscular strengthening and stability training). The conceptual thinking behind the design of this rocker bottom shoe is that the smaller intrinsic muscles that act closer to the axis of rotation are strengthened and can react faster during perturbations or destabilizing conditions. By using these smaller muscles with lines of actions closer to the axis, joint loading may be reduced. Their use as training, as opposed to clinical devices, offers quite different biomechanical results (e.g., gait mechanics, stability, plantar pressures, etc.) from more traditional rocker bottom shoes [7,64]. However, their efficacy as a true “fitness shoe” may be somewhat limited. The prescription of different rocker bottom shoes depends on the pathology to be addressed. A forefoot rocker incorporates a rocker, or angle to the sole of the shoe underneath or just behind the metatarsal heads. This is an effective design for reducing peak plantar pressures at the ball of the foot and reducing toe and forefoot sagittal plane motion. It may be used to control pain related to hallux limitus. A full foot (heel to toe) rocker is designed to limit ankle and midfoot motion and may sometimes be incorporated into a stabilizing boot after surgical treatments or injuries to limit range of motion during recovery. Full foot rockers are beneficial for pathologies such as ankle arthritis. Both the geometry of the rocker bottom and the length of the rocker may potentially affect energy expenditure and other biomechanical measures during walking [65].

41.8

Variations on materials

AFO materials vary depending on the needs of the patient (and reimbursement policy) and may include thermosetting materials, thermoplastics, composite type materials, fabric strapping, and/or carbon fiber. Both conventional and advanced AFOs may use a range of different materials but carbon fiber materials are more commonly found in advanced AFOs. Recent studies on advanced AFOs comprised of carbon fiber found that they improved gait (e.g., walking speed and overall performance) across a variety of patient populations from cerebral palsy [66] to stroke [67] to limb salvage [68]. When comparing carbon fiber advanced AFO designs to thermoplastic conventional designs, Patzkowski et al. [68] found that the carbon fiber design of the Intrepid Dynamic Exoskeletal Orthosis offered greater functional and performance benefits for military patients with lower limb salvage. Conventional AFOs are more often thermoplastic or thermoset construction, which may be simpler to manufacture than many other designs but provide less medial-lateral stability. These AFOs can be manufactured without an articulating joint, to provide more sagittal plane control or with an ankle joint to allow greater sagittal plane motion. In a systematic review on patients who had experienced a stroke, AFOs improved walking velocity regardless of the type of material with which the device was made [61].

41.9

New designs

New designs and applications for AFOs and rocker bottom shoes often show up clinically or in the literature. Patient needs often drive clinician creativity for finding new solutions whether it be from the design of the device, the material properties, or motorized joints. As an example, active AFO designs are relatively new in research and are being designed and developed to address the deficiencies of different clinical populations. These AFO designs use motors, pumps, and/or actuators to augment human function and provide joint assistance [69,70]. Addressing the lack of ankle push-off power is one of the more common—and highly desirable—design features for a wide variety of patient populations. 3D printing has also become an option for AFOs. While 3D printing techniques, such as selective laser sintering, may not yet be mainstream, some studies have shown it to be a viable option for fabricating different parts [71,72]. They allow for the incorporation of complex features without the added burden on manufacturing cost and time. They also provide opportunities to systematically vary certain design features such as weight, stiffness, or bending axis to explore new options for patients [71,73,74]. Overall, 3D printed AFOs—or AFO components—would be desirable for clinicians and patients as they offer the opportunity for greater degrees of customization. Several hurdles, including durability of materials, must be overcome for this to become a more mainstream option.

41.10 Sport applications Studies on AFOs and rocker bottom shoes have long been confined to pathological gait. However, sport applications have been presented recently in the literature. Some advanced AFOs were originally designed using many of the same

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mechanical principles as running-specific prosthetic feet. A deformable posterior strut absorbs and returns mechanical energy to beneficially affect performance. Individuals with limb salvage procedures, for example, have returned to activities such as running, jumping, plyometrics, sports participation, and military deployment [68,75 78]. An advanced AFO with a proximal patellar tendon-bearing ground reaction cuff connected by a carbon fiber posterior strut has improved performance on high impact and agility-based tasks compared to conventional AFOs [68]. Rocker bottom shoes have also transitioned into the athletic arenas in past decades. Specifically, rockers have been incorporated into running shoe designs for performance and mitigation of injury. Because high loads on the forefoot region may lead to overuse running injuries (e.g., metatarsal stress fractures or metatarsalgia) [79], rocker bottom shoes’ ability to reduce forefoot plantar pressures [63] may play a role in injury prevention. A proximally placed rocker shifts the ground reaction force vector posteriorly and may reduce the load on the Achilles tendon to potentially benefit runners who suffer from the common overuse injury, Achilles tendinopathy [80]. In a separate study, a rocker bottom shoe (e.g., Masai Barefoot Technologies) reduced sagittal plane ankle motion and the internal plantarflexor moment at push-off, while not significantly altering knee and hip mechanics [81]. These results indicate that this footwear intervention may be effective for runners returning to sport after Achilles tendinopathy.

41.11 Areas of future research AFOs and rocker bottom shoes can be effective tools for changing gait mechanics for therapeutic and performance benefits. However, the levels of scientific evidence currently available are often insufficient to adequately inform the prescription of AFOs and rocker bottom shoes for a variety of clinical populations. While many questions have been addressed, more remain unanswered and the long-term effects of using these devices, their benefits and limitations, the needs and preferences of the end user, and the optimal time to use them are still in question. As new designs and materials become available, the research behind these devices will expand to demonstrate their efficacy for certain patient populations. Nevertheless, they remain relatively inexpensive and low-risk options to try for a wide variety of pathologies.

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Arch Phys Med Rehabil 2004;85(1):81 6. [64] Stewart L, Gibson JN, Thomson CE. In-shoe pressure distribution in “unstable” (MBT) shoes and flat-bottomed training shoes: a comparative study. Gait Posture 2007;25(4):648 51. [65] Ruina A, Bertram JE, Srinivasan M. A collisional model of the energetic cost of support work qualitatively explains leg sequencing in walking and galloping, pseudo-elastic leg behavior in running and the walk-to-run transition. J Theor Biol 2005;237(2):170 92. [66] Bartonek A, Eriksson M, Gutierrez-Farewik EM. A new carbon fibre spring orthosis for children with plantarflexor weakness. Gait Posture 2007;25(4):652 6. [67] Danielsson A, Sunnerhagen KS. Energy expenditure in stroke subjects walking with a carbon composite ankle foot orthosis. J Rehab Med 2004;36(4):165 8. [68] Patzkowski JC, Blanck RV, Owens JG, Wilken JM, Kirk KL, Wenke JC, et al. Comparative effect of orthosis design on functional performance. J Bone Jt Surg Am 2012;94(6):507 15. [69] Gordon KE, Sawicki GS, Ferris DP. Mechanical performance of artificial pneumatic muscles to power an ankle-foot orthosis. J Biomech 2006;39(10):1832 41. [70] Blaya JA, Herr H. Adaptive control of a variable-impedance ankle-foot orthosis to assist drop-foot gait. IEEE Trans Neural Syst Rehabil Eng 2004;12(1):24 31. [71] Faustini MC, Neptune RR, Crawford RH, Stanhope SJ. Manufacture of Passive Dynamic ankle-foot orthoses using selective laser sintering. IEEE Trans Biomed Eng 2008;55(2 Pt 1):784 90. [72] Harper NG, Russell EM, Wilken JM, Neptune RR. Selective laser sintered vs carbon fiber passive-dynamic ankle-foot orthoses: a comparison of patient walking performance. J Biomech Eng 2014;136(9). [73] Ranz EC, Russell Esposito E, Wilken JM, Neptune RR. The influence of passive-dynamic ankle-foot orthosis bending axis on gait performance in individuals with lower-limb impairments. Clin Biomech 2016. [74] Harper NG, Russell Esposito E, Wilken JM, Neptune RR. The influence of ankle-foot orthosis stiffness on walking performance in individuals with lower-limb impairments. Clin Biomech (Bristol, Avon) 2014;29(8):877 84. [75] Patzkowski JC, Owens JG, Blanck RV, Kirk KL, Hsu JR, Skeletal Trauma Research C. Deployment after limb salvage for high-energy lowerextremity trauma. J Trauma Acute Care Surg 2012;73(2 Suppl. 1):S112 15. [76] Owens JG, Blair JA, Patzkowski JC, Blanck RV, Hsu JR, Skeletal Trauma Research C. Return to running and sports participation after limb salvage. J Trauma 2011;71(1 Suppl.):S120 4. [77] Blair JA, Patzkowski JC, Blanck RV, Owens JG, Hsu JR, Skeletal Trauma Research C. Return to duty after integrated orthotic and rehabilitation initiative. J Orthop Trauma 2014;28(4):e70 4. [78] Bedigrew KM, Patzkowski JC, Wilken JM, Owens JG, Blanck RV, Stinner DJ, et al. Can an integrated orthotic and rehabilitation program decrease pain and improve function after lower extremity trauma? Clin Orthop Relat Res 2014.

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[79] Hockenbury RT. Forefoot problems in athletes. Med Sci Sports Exerc 1999;31(7 Suppl):S448 58. [80] Sobhani S, Zwerver J, van den Heuvel E, Postema K, Dekker R, Hijmans JM. Rocker shoes reduce Achilles tendon load in running and walking in patients with chronic Achilles tendinopathy. J Sci Med Sport 2015;18(2):133 8. [81] Boyer KA, Andriacchi TP. Changes in running kinematics and kinetics in response to a rockered shoe intervention. Clin Biomech (Bristol, Avon) 2009;24(10):872 6.

Chapter 42

Diabetic Footwear Sicco A. Bus Department of Rehabilitation Medicine, Amsterdam UMC, University of Amsterdam, Amsterdam Movement Sciences, Amsterdam, The Netherlands

Abstract People with diabetic foot disease are commonly prescribed with special footwear or other devices to help prevent or treat a foot ulcer. The main mechanism here is the redistribution of mechanical pressure on the foot so to offload specific high-risk or ulcer regions. The most effective way to offload a plantar foot ulcer is by using a total contact cast or a knee-high walker; ankle-high devices are commonly used for offloading but are less effective. Footwear to help prevent a (recurrent) foot ulcer is often custom-made and consists of pressure relieving elements such as a rocker outsole, custom-made insert, metatarsal pad or bar, and cushioning insole top layer. In-shoe plantar pressure measurement proves to be a valuable tool to improve the pressure-relieving properties of preventative footwear, and when such pressure-improved footwear is adequately worn, a significant number of recurrent foot ulcers can be prevented. Guidelines for what entails improved pressure relief have been developed, but ways to improve footwear adherence are not well known; educational techniques and specific offloading shoes for use indoors may contribute. Offloading is one of the cornerstones of treatment of diabetic foot disease and should always be considered in high-risk patients or those with active plantar foot ulcers.

42.1

Introduction

People with diabetic foot disease are commonly prescribed with special footwear or other devices for treating a foot ulcer or to help prevent such an ulcer from developing. One of the main mechanisms by which these goals are achieved is redistribution of mechanical pressure applied to the foot. After discussing these biomechanical mechanisms in more detail, the different footwear and other devices used for ulcer healing and prevention will be discussed and the importance of adherence will be highlighted.

42.2

Foot biomechanics and offloading

Of people with diabetes mellitus, approximately 50% will eventually have loss of protective sensation (LOPS) in their feet due to peripheral neuropathy developing as a complication of the disease [1]. This LOPS is sufficient to allow these patients to injure their feet without notice [2]. In the presence of LOPS, elevated plantar foot pressure during ambulation is a causative factor in the development of foot ulcers and also in the recurrence thereof [3,4]. These elevated plantar foot pressures are generally caused by changes in foot structure, such as deformity, limited joint mobility, and loss of fat pad quality [5]. Reducing these mechanical pressures under the foot is called offloading. If the foot is not offloaded adequately, ulcers may develop or current ulcers may not heal. Offloading is considered one of the cornerstones of treatment for and against plantar foot ulcers in people with diabetes [6]. After an ul‘/cer is healed, the risk of recurrence is high, up to 40% in the first 12 months, and 60% 65% after 3 years [2]. This shows the need for continuous offloading in these high-risk patients. The offloading through footwear or other devices can be measured inside the shoe or device with a plantar pressure measurement system. This is an insole with an array of sensors that is inserted inside the shoe or device and measures the plantar pressure distribution. Pressure measurement is now becoming more widespread in use in clinical practice, Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00014-7 © 2023 Elsevier Inc. All rights reserved.

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mainly through the effect of several studies demonstrating the benefit of such measurements and that are discussed below. A detailed discussion of mechanical pressure in diabetic foot disease can be found in other sections of this book and elsewhere [7].

42.3

The biomechanical effect of diabetic footwear and offloading devices

For the offloading treatment of plantar foot ulcers, different devices and shoes are used. The most effective way to offload the plantar foot is by using a total contact cast (TCC) or a knee-high walker [8,9]. These devices can reduce forefoot peak pressure by up to 80% 90% compared to standard nontherapeutic shoes (Fig. 42.1) [8]. This effect is achieved by redistribution of forces not only across the plantar foot surface, but also to the lower leg through the cone shape of the device wall, in addition to limiting ankle motion [10,11]. Offloading devices that only extend to or just above the ankle, such as cast shoes and forefoot offloading shoes, are generally less effective than a TCC in offloading the foot, with approximately 40% 60% peak pressure relief found compared to a standard control shoe condition [8,12,13]. Therapeutic footwear deals with shoes and insoles (also called inserts or orthoses) that aim for some therapeutic effect, and in people with diabetic foot diseases this is mostly the prevention of foot ulcers through the provision of sufficient space and pressure relief at high-risk areas. Such footwear is built up from different pressure-relieving design elements and these different elements have also been studied for their capacity to offload the foot [14]. A rocker-

FIGURE 42.1 Peak plantar pressure distribution shown for the left and right foot as measured inside custom-made shoes (left) and in a total contact cast (right) in the same patient who had a foot ulcer on the plantar hallux. Note the effective pressure relief over the entire plantar surface by use of the total contact cast.

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bottom outsole configuration is very effective in relieving plantar pressure in the forefoot, with up to 52% of pressure relief found compared to a shoe that does not have a rocker outsole [15 17]. Custom-made insoles are effective for relieving peak pressure compared to standard flat insoles [18]. Furthermore, specific design elements of an insole such as metatarsal pads and bars, and medial arch supports have shown to be effective in relieving pressure in regions where foot ulcers in people with diabetes often occur [19,20]. Additionally, the choice of top layer of an insole results in significant relief of peak pressure over the entire plantar surface and more local modifications such as removing and softening insole materials at at-risk regions are also effective [20]. It should be noted that while such insole design elements can effectively offload plantar foot regions, the specific form, materials used and placement of these metatarsal pads and bars, and medial arch supports is critical in the final pressure outcome; and incorrect placement can lead to pressure increases instead of decreases [21]. The use of in-shoe plantar pressure measurement is valuable to optimize placement of these insole elements [22,23].

42.4

Footwear and offloading for ulcer healing

In healing plantar foot ulcers, specific offloading devices for temporary use are most commonly prescribed (Fig. 42.2), in preference to footwear modalities that have lower offloading efficacy. Most evidence on the use of offloading devices is for treating neuropathic plantar forefoot and midfoot ulcers that are not complicated by infection or ischemia [9]. These represent about 25% of foot ulcers seen in multidisciplinary expert diabetic foot clinics [24]. While the direct association between the use of offloading and healing of a foot ulcer has been studied to only a limited extent [25,26], indirect evidence using capacity to offload and efficacy to heal demonstrates this close relationship [8]. Ulcer healing is mostly expressed as proportion of patients healed in a given time (e.g., 12 or 20 weeks) and time to complete healing. Multiple recent and well-conducted meta-analyses and systematic reviews on offloading treatment of plantar foot ulcers consistently show that nonremovable knee-high offloading is more effective than any form of knee-high or

FIGURE 42.2 Offloading devices used for the treatment of plantar foot ulcers in people with diabetes. (A) removable total contact cast; (B) anklehigh cast shoe; (C) removable knee-high walker; (D) removable knee-high walker; (E) total contact cast; (F) removable walker rendered irremovable; (G) ankle-high cast shoe; (H) vacuum-assisted tibia-high walker; (I) felted foam; (J) forefoot offloading shoe; (K) custom-made insole; (L) temporary custom-made shoe; (M) custom-made shoe.

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ankle-high removable offloading [9,27 30]. And within the category of knee-high nonremovable offloading, the TCC and walker rendered nonremovable show to have comparable efficacy in ulcer healing [9]. These findings suggest that it is no longer only the TCC, but rather offloading in a nonremovable knee-high device, that is considered the “gold standard” in the treatment of plantar forefoot ulcers in patients with diabetes. Based on limited evidence available, these knee-high nonremovable devices have now also been indicated as the first-choice treatment modality for healing plantar midfoot and hindfoot ulcers [31]. Despite the clear hierarchy in offloading effectiveness, ankle-high devices are commonly used for offloading treatment of plantar foot ulcers and include forefoot offloading shoes, cast shoes, and custom-made temporary shoes (Fig. 42.2). These footwear modalities may be effective in healing neuropathic forefoot ulcers, but their efficacy has often only been assessed in retrospective studies, and not in high-quality randomized controlled trials [9,32 35]; confirmation in such trials is required, mainly because of their common use in clinical practice.

42.5

Diabetic footwear for ulcer prevention

Therapeutic footwear had been extensively studied for its ability to prevent the occurrence of foot ulcers (Fig. 42.3). Most of these studies have focused on the prevention of ulcer recurrence as patients are normally only prescribed with such footwear when a foot problem like an ulcer has already occurred. Very few studies exist on the prevention of a first ever foot ulcer, which is also understandable from the point of view that this requires a much larger study population due to the lower event rate of ulceration than with ulcer recurrence. Several prospective studies from before 2013 have shown a beneficial effect of the use of therapeutic footwear compared to standard footwear in preventing ulcer

FIGURE 42.3 Examples of therapeutic shoes and insoles that are prescribed to people with diabetes who are at high risk of developing a foot ulcer.

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recurrence [36]. One large RCT showed no effect [36,37]. These contrasting results can be attributed to a wide diversity of intervention and control conditions tested and the lack of information about the offloading effect of the footwear used, complicating the comparison of studies in this area [38]. Only quite recently, two multicenter RCTs have further improved our understanding of the role of offloading in therapeutic footwear for preventing foot ulcer recurrence [39,40]. In one trial from our group, in-shoe plantar pressure analysis was used as a tool to guide modifications to custom-made footwear that was prescribed and delivered to patients based on the expertise and skills of the clinical team, not using scientific knowledge on footwear efficacy in a systematic way. This approach significantly improved the pressure-relieving properties of the custom-made footwear, with an average 30% peak pressure relief, but it showed only a nonsignificant 11% reduction in incidence of ulcer recurrence after 18 months compared to custom-made footwear that did not undergo such modifications and improvement based on pressure guidance [39]. However, when only the patients who, with objective measures [41], were adherent to wearing their prescribed footwear were analyzed, pressure-improved footwear showed a significant 46% lower plantar foot ulcer incidence rate compared to the normal shoes that did not undergo this improvement. In the other trial, a multicenter RCT from the United States, custom-made insoles that were designed and manufactured based on barefoot plantar pressure and 3D foot shape data and worn in extra-depth shoes were compared to traditional shape-based custom-made insoles for efficacy to prevent plantar forefoot ulcer recurrence over 15 months. The pressure-based insole reduced risk of ulcer recurrence by 63% [40]. These data-driven footwear concepts demonstrate that it is the combination of adequate pressure relief through pressure-based design and adequate adherence to wearing the footwear that leads to the best clinical outcomes. What adequate pressure relief means quantitatively is not exactly known, and will differ depending on the type of in-shoe plantar pressure measurement system used. But we have some useful indications from comparative and prospective studies. One study examined patients who had remained healed after plantar ulceration and found a mean pressure of approximately 200 kPa at the prior ulcer site [42]. Data from the above-mentioned trial showed that peak pressures .200 kPa could be effectively reduced, but when pressures were ,200 kPa, further adaptation of the footwear proved futile [23]. Additionally, a risk factor analysis of pressure-related plantar ulcer recurrence in the same trial showed that when peak pressure at at-risk regions is below 200 kPa and adherence is above 80%, the risk of plantar foot ulcer recurrence is 60% lower compared to when these conditions are not met [4]. Even though a pressure threshold for foot ulceration is likely unique to each individual, we now have indications that the 200kPa value can serve as a useful target for plantar offloading in footwear prescription [23]. This is under the condition that the pressures are measured with a validated and calibrated system with a spatial resolution of at least 1 sensor per 2 cm2 [43]. The recommendation for demonstrated plantar pressure relief is now part of international guidelines on the prevention of foot ulcers from the International Working Group on the Diabetic Foot (IWGDF) [43]. This illustrates that diabetic footwear provision for ulcer prevention is moving from a skills and experience-based method to a more systematic, scientific and data-driven approach. The effects of such an approach have been studied recently by comparing several data-driven scientific-based footwear design methods on plantar pressure relief and show that more effective pressure relief and less need for modification of footwear after in-shoe pressure measurement are provided by these data-driven footwear designs [44].

42.6

Footwear and offloading adherence

Regardless of their pressure-relieving capacity, diabetic footwear and offloading devices can only be effective when they are worn by the patient. Therefore, these shoes and devices must be judged both by their capacity to relief peak pressure and by the patient’s adherence to the treatment. Studies show that ulcer treatment with a removable offloading device may be complicated by nonadherence of the patient. Patients with an active foot ulcer may use their prescribed removable knee-high walker for an average of only 29% of their total daily number of steps [45]. And of patients who are at high risk of developing a foot ulcer, considering they have peripheral neuropathy, a foot ulcer history, and foot deformity, only 50% may wear their prescribed custom-made shoes for more than 80% of the steps they take; nonadherence is particularly a problem when patients are inside their home where they take more steps than when outside [46]. This contributes to the lower effectiveness of footwear and removable offloading devices and stresses the importance of continued pressure relief to promote healing and prevent foot ulcers [14,26]. Despite the important role of adherence, studies on how to improve adherence with wearing diabetic footwear and removable offloading devices are scarce [14,46]. One small explorative study on the effect of motivational interviewing as behavioral intervention to improve offloading shows variable effects between patients and some effect only in the short term [47]. A very recent study on the effect of specific custom-made shoes designed for indoor use showed significantly improved adherence indoors, both short and long-term after prescribing such footwear next to the custom-made

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shoes that the patient already has [48]. Patients seem to be satisfied with having such a shoe for special use indoors, as these shoes are much easier to don and doff, are less heavy and warm, and not dirty from outside use, while still having comparable offloading properties as the custom-made shoes they already have [49]. More development of methods and tools that can improve adherence and studies testing these methods is needed.

42.7

Other considerations

The evidence shows that neuropathic plantar forefoot ulcers that are not complicated by infection or ischemia can be healed in approximately 6 8 weeks (and 90% in 12 weeks time), if adequate offloading is used, primarily through the use of nonremovable knee-high offloading devices [9]. However, several studies and surveys show that the devices for which most evidence exists are underused in clinical practice, including only 2% use of the TCC in one survey; shoes and felted foam, for which hardly any evidence exists, are most commonly used for treating plantar foot ulcers [50 52]. This gap between recommendations and use in clinical practice needs to be bridged [8]. Approximately 75% of ulcer in specialized foot clinics are complicated by infection and ischemia [24,53]. These ulcers take longer to heal than neuropathic foot ulcers and treatment requires a multidisciplinary approach in which primarily the infection and ischemia are under control and secondarily offloading treatment is initiated [54,55]. Offloading is still important in such complex wounds because of the neuropathy present and the enhanced risk of limb loss in these patients. The most recent IWGDF guideline on offloading has now included several recommendations on the use of offloading in the treatment of more complex plantar foot ulcers [31]. In quantifying offloading, we normally refer to the normal component of pressure, not to the shear component, as this cannot be measured with the currently available systems. In ulcer development and healing, shear plays a contributing role. For example, the TCC immobilizes the ankle joint which is thought to limit movement of the foot inside the device and, with that, reduce shear. And hyperkeratotic tissue or callus, under which foot ulcer often occurs, is considered to be the result of shear acting on the foot [2]. Given the lack of available shear measuring options, foot temperature has been opted as a surrogate for shear [56], which, if shown to be associated would open up options for shear assessment inside footwear and offloading devices. Clearly, the topic of shear requires more scientific investigation. Despite the value of using plantar pressure measurement in the design and evaluation of diabetic footwear and offloading devices, it is not commonly used in clinical practice. A major advance would be the requirement that measurable and effective pressure reduction should result from all prescribed interventions [8]. Requirements for demonstrated efficacy in pressure relief for prescribed custom-made shoes have been introduced in Germany, and insurance companies in the Netherlands are now also considering to make this part of the requirement for reimbursement of custom-made shoes for diabetic patients. Such measurements may not be possible at every treatment location because of available resources, but a cost-analysis of our footwear trial does show the cost-benefit of implementing such pressure measurements in clinical footwear practice [39]. Otherwise, specialized centers or footwear companies should consider adding the measurement of plantar pressure to their prescription approach, so to provide state-of-the-art knowledge and tools to diabetic footwear practice.

42.8

Future research

Several aspects related to the offloading treatment of people with diabetic foot disease require further research. More research and development should be done in measuring shear stress as a likely important component of ulcer development in people with diabetes and neuropathy. Shear stress is technically difficult to assess, and foot temperature has been opted as a surrogate for shear. More research in this association is needed. More development is also needed in methods and tools that can improve adherence to wearing prescribed footwear, and studies on their efficacy are urgently needed. More research in the offloading efficacy of more complex foot ulcers that are infected or ischemic is needed, due to the enhanced risk of limb loss and because these ulcers represent the majority seen nowadays in diabetic foot practice. And finally, future research should focus on defining successes and failures in implementing the measurement of in-shoe plantar pressure in footwear prescription by hospitals and footwear companies to provide state-of-the-art knowledge and tools to diabetic footwear practice.

42.9

Conclusions

This chapter discusses diabetic footwear and offloading treatment for people with diabetes who either have an active foot ulcer or are at high risk of developing one. The “gold standard” treatment for neuropathic plantar foot ulcers is a nonremovable knee-high offloading device, which is more effective in offloading and in healing foot ulcers compared

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to any removable knee-high or ankle-high device. Its limited use, however, in clinical practice is a problem and needs to be addressed. Diabetic footwear with a demonstrated pressure-relieving effect has shown to be effective in reducing risk of plantar foot ulcer recurrence in people with diabetes, under the condition that the footwear is worn. The recommendation for such pressure-relieving footwear is now integrated in international guidelines on the diabetic foot. Custom-made footwear design protocols and the implementation of plantar pressure measurement help in translating these guidelines to clinical footwear practice. With that, diabetic footwear design and provision moves from a traditional field to a modern field that uses a systematic, science-based, and data-driven approach that clinically benefits the high-risk patient.

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[26] Crews RT, Shen BJ, Campbell L, Lamont PJ, Boulton AJ, Peyrot M, et al. Role and determinants of adherence to off-loading in diabetic foot ulcer healing: a prospective investigation. Diabetes Care 2016;39(8):1371 7. [27] Lewis J, Lipp A. Pressure-relieving interventions for treating diabetic foot ulcers. Cochrane Database Syst Rev 2013;1:CD002302. [28] Morona JK, Buckley ES, Jones S, Reddin EA, Merlin TL. Comparison of the clinical effectiveness of different off-loading devices for the treatment of neuropathic foot ulcers in patients with diabetes: a systematic review and meta-analysis. Diabetes Metab Res Rev 2013;29(3):183 93. [29] Elraiyah T, Prutsky G, Domecq JP, Tsapas A, Nabhan M, Frykberg RG, et al. A systematic review and meta-analysis of off-loading methods for diabetic foot ulcers. J Vasc Surg 2016;63(2 Suppl.):59S 68S e2. [30] Fibreglass Total Contact Casting. Removable cast walkers, and irremovable cast walkers to treat diabetic neuropathic foot ulcers: a health technology assessment. Ont Health Technol Assess Ser 2017;17(12):1 124. [31] Bus SA, Armstrong DG, Gooday C, Jarl G, Caravaggi C, Viswanathan V, et al. Guidelines on offloading foot ulcers in persons with diabetes (IWGDF 2019 update). Diabetes Metab Res Rev 2020;36(Suppl. 1):e3274. [32] Hissink RJ, Manning HA, van Baal JG. The MABAL shoe, an alternative method in contact casting for the treatment of neuropathic diabetic foot ulcers. Foot Ankle Int 2000;21(4):320 3. [33] Dumont IJ, Lepeut MS, Tsirtsikolou DM, Popielarz SM, Cordonnier MM, Fayard AJ, et al. A proof-of-concept study of the effectiveness of a removable device for offloading in patients with neuropathic ulceration of the foot: the Ransart boot. Diabet Med 2009;26(8):778 82. [34] Van De Weg FB, Van Der Windt DA, Vahl AC. Wound healing: total contact cast vs. custom-made temporary footwear for patients with diabetic foot ulceration. Prosthet Orthot Int 2008;32(1):3 11. [35] Bus SA, van Netten JJ, Lavery LA, Monteiro-Soares M, Rasmussen A, Jubiz Y, et al. IWGDF guidance on the prevention of foot ulcers in atrisk patients with diabetes. Diabetes Metab Res Rev 2016;32(Suppl. 1):16 24. [36] van Netten JJ, Raspovic A, Lavery LA, Monteiro-Soares M, Rasmussen A, Sacco ICN, et al. Prevention of foot ulcers in the at-risk patient with diabetes: a systematic review. Diabetes Metab Res Rev 2020;36(Suppl. 1):e3270. [37] Reiber GE, Smith DG, Wallace C, Sullivan K, Hayes S, Vath C, et al. Effect of therapeutic footwear on foot reulceration in patients with diabetes: a randomized controlled trial. JAMA 2002;287(19):2552 8. [38] Cavanagh PR, Bus SA. Off-loading the diabetic foot for ulcer prevention and healing. J Am Podiatr Med Assoc 2010;100(5):360 8. [39] Bus SA, Waaijman R, Arts M, de HM, Busch-Westbroek T, Van BJ, et al. Effect of custom-made footwear on foot ulcer recurrence in diabetes: a multicenter randomized controlled trial. Diabetes Care 2013;36(12):4109 16. [40] Ulbrecht JS, Hurley T, Mauger DT, Cavanagh PR. Prevention of recurrent foot ulcers with plantar pressure-based in-shoe orthoses: the CareFUL prevention multicenter randomized controlled trial. Diabetes Care 2014;37(7):1982 9. [41] Bus SA, Waaijman R, Nollet F. New monitoring technology to objectively assess adherence to prescribed footwear and assistive devices during ambulatory activity. Arch Phys Med Rehabil 2012;93(11):2075 9. [42] Owings TM, Apelqvist J, Stenstrom A, Becker M, Bus SA, Kalpen A, et al. Plantar pressures in diabetic patients with foot ulcers which have remained healed. Diabet Med 2009;26(11):1141 6. [43] Bus SA, Lavery LA, Monteiro-Soares M, Rasmussen A, Raspovic A, Sacco ICN, et al. Guidelines on the prevention of foot ulcers in persons with diabetes (IWGDF 2019 update). Diabetes Metab Res Rev 2020;36(Suppl. 1):e3269. [44] Zwaferink JB, Custers W, Paardekoper I, Berendsen H, Bus SA. Optimizing footwear for the diabetic foot: Data-driven custom-made footwear concepts and their effect on pressure relief to prevent diabetic foot ulceration. Plos One 2020;15(4):e0224010. [45] Armstrong DG, Lavery LA, Kimbriel HR, Nixon BP, Boulton AJ. Activity patterns of patients with diabetic foot ulceration: patients with active ulceration may not adhere to a standard pressure off-loading regimen. Diabetes Care 2003;26(9):2595 7. [46] Waaijman R, Keukenkamp R, de haart M, Polomski WP, Nollet F, Bus SA. Adherence to wearing prescription custom-made footwear in patients with diabetes at high risk for plantar foot ulceration. Diabetes Care 2013;36(6):1613 18. [47] Keukenkamp R, Merkx MJ, Busch-Westbroek TE, Bus SA. An explorative study on the efficacy and feasibility of the use of motivational interviewing to improve footwear adherence in persons with diabetes at high risk for foot ulceration. J Am Podiatr Med Assoc 2018;108(2):90 9. [48] Keukenkamp R, van Netten JJ, Busch-Westbroek TE, Bus SA. Custom-made footwear designed for indoor use increases short-term and longterm adherence in people with diabetes at high ulcer risk. BMJ Open Diabetes Res Care 2022;10(1):e002593. [49] Keukenkamp R, van Netten JJ, Busch-Westbroek TE, Nollet F, Bus SA. Users’ needs and expectations and the design of a new custom-made indoor footwear solution for people with diabetes at risk of foot ulceration. Disabil Rehabil 2021;1 8. [50] Wu SC, Jensen JL, Weber AK, Robinson DE, Armstrong DG. Use of pressure offloading devices in diabetic foot ulcers: do we practice what we preach? Diabetes Care 2008;31(11):2118 19. [51] Prompers L, Huijberts M, Apelqvist J, Jude E, Piaggesi A, Bakker K, et al. Delivery of care to diabetic patients with foot ulcers in daily practice: results of the Eurodiale Study, a prospective cohort study. Diabet Med 2008;25(6):700 7. [52] Raspovic A, Landorf KB. A survey of offloading practices for diabetes-related plantar neuropathic foot ulcers. J Foot Ankle Res 2014;7:35. [53] Cavanagh PR, Lipsky BA, Bradbury AW, Botek G. Treatment for diabetic foot ulcers. Lancet. 2005;366(9498):1725 35. [54] Nabuurs-Franssen MH, Sleegers R, Huijberts MS, Wijnen W, Sanders AP, Walenkamp G, et al. Total contact casting of the diabetic foot in daily practice: a prospective follow-up study. Diabetes Care 2005;28(2):243 7. [55] Bus SA, Armstrong DG, van Deursen RW, Lewis JE, Caravaggi CF, Cavanagh PR, et al. IWGDF guidance on footwear and offloading interventions to prevent and heal foot ulcers in patients with diabetes. Diabetes Metab Res Rev 2016;32(Suppl. 1):25 36. [56] Yavuz M, Brem RW, Davis BL, Patel J, Osbourne A, Matassini MR, et al. Temperature as a predictive tool for plantar triaxial loading. J Biomech 2014;47(15):3767 70.

Chapter 43

Reconstructions for Adult-acquired Flatfoot Deformity Matthew S. Conti, Jonathan H. Garfinkel and Scott J. Ellis Hospital for Special Surgery (HSS), New York, NY, United States

Abstract Adult-acquired flatfoot deformity ranges in severity from inflammation of the posterior tibial tendon to rigid deformities and talar tilt. For patients with mild symptoms, conservative management may be attempted; however, patients with persistent pain and bony deformities are commonly treated with a flatfoot reconstruction, which can improve symptoms and prevent progression of the disease. The multiple deformities associated with the adult-acquired flatfoot can be explained by the failure of the ligaments, tendons, and bony structures that support the talus, which leads to peritalar subluxation. This chapter reviews how the failure of each of these components alters the biomechanics of the foot, contributes to peritalar subluxation, and how each deformity can be reconstructed. The chapter is organized by deformity and includes hindfoot valgus, forefoot external rotation, sag at the talonavicular joint, failure of the posterior tibial tendon, gastrocnemius and Achilles tendon tightness, and medial arch eversion. The chapter concludes with special considerations when reconstructing the adultacquired flatfoot and future biomechanical studies that can help surgeons to better understand the associated deformities.

43.1

Introduction

Adult-acquired flatfoot deformity (AAFD) is characterized by multiple deformities including hindfoot valgus, forefoot external rotation, sag at the talonavicular joint, and medial arch eversion. It is more common in women, and the prevalence of symptomatic AAFD in females over the age of 40 is estimated to be 3% with a peak incidence at age 55 [1,2]. The cause of the condition is multifactorial and associated with preexisting adolescent flatfoot, obesity, and factors affecting the vascularity of the posterior tibial tendon [2,3]. Although first described as posterior tibial tendon dysfunction, the deformities associated with the adult flatfoot cannot be attributed to insufficiency of a single structure but are better explained by the failure of the ligaments, tendons, and bony structures that support the talus [2,4]. As these structures fail, there is the development of progressive peritalar subluxation, which can be characterized from stage I to IV according to the classification system initially developed for posterior tibial tendon insufficiency [2,5 7]. In stage I, there is tenosynovitis (tendon sheath inflammation) of the posterior tibial tendon with a radiographically normal foot. Stage II is divided into stage IIA and stage IIB. Stage IIA has collapse of the medial longitudinal arch, a flexible hindfoot deformity, and inability to perform a single heel rise but no significant external rotation of the forefoot; whereas, stage IIB has the same findings as stage IIA but with an external rotation deformity of the forefoot described as greater than 30% 40% of talonavicular uncoverage on a standing AP radiograph [2]. In stage III, the hindfoot and forefoot deformities are rigid, and radiographs may begin to show signs of subtalar arthritis. In the final stage, stage IV, there is rigid deformity and insufficiency of the deltoid ligament with talar tilt seen on the mortise view of the ankle. This progression is a result of failures in one or more of the structures supporting the talus. Early surgical reconstructions for AAFD focused on the correction of either the hindfoot valgus or the forefoot external rotation deformities. In 1971, Koutsogiannis popularized the medializing calcaneal osteotomy (MCO) originally proposed by Gleich in 1893 to treat the flexible flatfoot deformity [8]. Following this, Evans proposed a lateral column lengthening (LCL) to treat calcaneovalgus [9]. Twenty-two years later, Pomeroy and Manoli introduced a reconstructive procedure for stage II disease that combined a flexor digitorum longus (FDL) transfer, MCO, and LCL Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00029-9 © 2023 Elsevier Inc. All rights reserved.

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to treat both the hindfoot valgus and forefoot external rotation deformities associated with AAFD [10]. Treatment of stage III AAFD typically involves the use of a double or triple arthrodesis [2,7]. Because stage IV disease involves the tibiotalar joint, management of stage IV disease requires addressing the ankle joint in addition to the planovalgus deformity [2]. This can be accomplished through a deltoid reconstruction, total ankle replacement, and/or ankle fusion in conjunction with the procedures employed in stage II disease if the deformity is flexible or through a combination of a double/triple arthrodesis and reconstructive procedures if the deformity is rigid [2]. In this chapter, we explain how the failure of ligaments, tendons, and bony structures leads to peritalar subluxation and changes in the biomechanics of the foot and ankle. We divide the chapter by deformity and then discuss the corresponding anatomy, biomechanics, pathoanatomy, and the potential surgical reconstruction options. We conclude the chapter with special considerations when reconstructing the adult-acquired flatfoot and future biomechanical studies that can help us to better understand its associated deformities.

43.2

Hindfoot valgus

The hindfoot valgus deformity occurs when the weightbearing axis of the tibia lies medial to the axis of the calcaneus. In addition to identifying the severity of the deformity, it is most important to assess the flexibility of the hindfoot as this will guide treatment. Patients with severe hindfoot valgus often cannot perform a single heel rise due to insufficiency of the posterior tibial tendon. When performing a double heel rise, the hindfoot will invert at terminal plantarflexion, which identifies the deformity as flexible; otherwise, if the hindfoot stays in valgus during the double heel rise, the deformity is rigid. In adolescents, tarsal coalition may cause rigid hindfoot valgus. Talocalcaneal and calcaneonavicular coalition may lead to limited subtalar motion resulting in rigid hindfoot valgus during toe-off.

43.2.1 Bony anatomy Radiographically, hindfoot valgus is best assessed on the hindfoot alignment view described by Saltzman and elKhoury (Fig. 43.1) [11]. To quantify the severity of the hindfoot valgus deformity, they established the hindfoot

FIGURE 43.1 Hindfoot alignment view radiograph in a normal patient (A). A normal hindfoot moment arm is approximately 3 mm of varus (4 mm of varus in this patient). A valgus hindfoot on the hindfoot alignment view radiograph in a patient with a flatfoot deformity (B). The hindfoot moment arm measures 30.5 mm in valgus.

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FIGURE 43.2 Talocalaneal impingement on a sagittal weightbearing CT scan (A). Patients with significant talocalcaneal impingement may require a lateral talar process debridement, lateral column lengthening, or subtalar fusion to alleviate their symptoms. The white arrowhead identifies talocrural impingement while the white arrow shows talocalcaneal impingement. A coronal slice of a weightbearing CT scan that illustrates calcaneofibular impingement (B). Patients with calcaneofibular impingement may require an adequate medializing calcaneal osteotomy or subtalar fusion to alleviate their symptoms. The white arrow in this image demonstrates the site of calcaneofibular impingement.

moment arm, which is the distance between the weight-bearing axis of the tibia to the most inferior point on the calcaneus [11]. Other authors have proposed additional measurements to quantify the severity of the hindfoot valgus deformity [12 16]. Lateral subluxation of the calcaneus beneath the talus results in hindfoot valgus in AAFD. Patients with stage II AAFD were found to have increased subtalar joint valgus compared with control patients [17]. This deformity is best appreciated on multiplanar weightbearing three-dimensional imaging, which better assesses hindfoot valgus, subtalar impingement, and calcaneofibular impingement than traditional nonweightbearing CT scans [17 19]. Understanding the severity of talocalcaneal and calcaneofibular impingement in patients with AAFD may help to determine operative treatment (Fig. 43.2). Patients with significant subtalar or talocalcaneal impingement may benefit from debridement of the lateral talar process or a LCL; whereas, performing an adequate MCO can alleviate calcaneofibular impingement [4]. In severe cases, these patients may benefit from a subtalar fusion. As the hindfoot moves into valgus, this affects the weight distribution across the foot. The heel typically provides static support for maintenance of the medial longitudinal arch of the foot, and lateral translation of the weightbearing axis of the calcaneus puts increased strain on the arch of the foot [20,21]. The deformity, therefore, increases the pressure on the soft tissue structures that help support the medial longitudinal arch, which leads to further attenuation of these ligaments and tendons ultimately resulting in the collapse of the arch. Additionally, the calcaneal tuberosity is the insertion site for the Achilles tendon. In normal feet, the Achilles tendon acts as an inverter of the heel. This pulls the heel into varus and locks the transverse tarsal joint during toe-off [21,22]. A valgus heel alignment changes the force vector of the Achillestendon and causes the Achilles to act as an everter of the heel [22]. This prevents the transverse tarsal joint from making the foot a rigid lever arm during toe-off and may lead to additional collapse of the medial longitudinal arch of the foot [22].

43.2.2 Ligament failure The hindfoot valgus deformity results primarily from insufficiency of the interosseous ligament that connects the talus and the calcaneus [4,23,24]. Although the posterior tibial tendon and plantar calcaneonavicular ligament (spring ligament) are thought to be the primary soft tissue structures that fail in AAFD, early stages (i.e., stage II) present with a valgus hindfoot that likely occurs only in the setting of a mechanically compromised interosseous ligament [24,25]. The talocalcaneal interosseous ligament was the third most commonly involved ligament after the superomedial and inferior calcaneonavicular ligaments in an MRI study involving 31 patients with AAFD (Fig. 43.3) [25]. In patients with severe hindfoot valgus deformity, failure of the deltoid ligament may be present [26]. Weightbearing ankle radiographs demonstrate talar tilt as a consequence of lack of soft tissue support of the medial malleolus. Weightbearing CT scans may help to better understand the bony anatomy including subluxation of the subtalar and tibiotalar joints as well as impingement on the lateral side of the ankle. MRI is a useful tool for evaluating the condition of the deltoid ligament when subtle talar tilt is found on plain radiographs. Deltoid ligament integrity has important implications in the surgical management of hindfoot valgus.

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FIGURE 43.3 Degenerated spring ligament on coronal MRI. Direct coronal proton density weighted MR image through the hindfoot in the setting of moderately severe adult-acquired flatfoot deformity (AAFD) with hindfoot valgus demonstrates markedly degenerated, thickened, and stretched superomedial fibers of the spring ligament (white arrows).

43.2.3 Surgical reconstruction Surgical reconstruction of the flexible hindfoot valgus deformity typically involves the use of a MCO [22,27 39]. The osteotomy translates the posterior calcaneus and calcaneal tuberosity medially providing static support to the medial longitudinal arch as well as moving the insertion of the Achilles tendon medially to restore its function as an inverter of the heel during toe-off [22]. The MCO has also been shown to alleviate a significant amount of strain on the spring ligament and restore its function [36,40]. While biomechanical studies have demonstrated that approximately 1 cm of translation of the posterior calcaneal segment provides significant offloading of the arch [36], studies using patient-reported outcomes have demonstrated better results with individualized translation of the heel [20]. In general, surgeons should aim for a hindfoot moment arm of 0 5 mm in varus on postoperative hindfoot alignment view radiographs, which results in a clinically straight heel (Fig. 43.4) [20]. A previous study demonstrated a predictable change in the preoperative to postoperative hindfoot moment arm based on the amount of intraoperative medial translation of the heel during an MCO [41]. In addition to the MCO, a LCL also affects the hindfoot valgus deformity [42]. Although typically employed to correct the forefoot external rotation deformity, a study using 12 cadaveric specimens demonstrated that the LCL can affect up to 60% of the hindfoot valgus deformity [42]. As a consequence, overall correction of hindfoot valgus during a flatfoot reconstruction results from a combination of an MCO and LCL. This should be kept in mind to avoid overcorrecting the deformity, which may increase plantar pressures and discomfort in the lateral foot [22]. In patients with a rigid hindfoot deformity, a double or triple arthrodesis is typically required to adequately correct the deformity [2,7]. Subtalar arthritis or subtalar impingement, both of which are best evaluated on weightbearing CT scans, are also indications for fusion of the subtalar joint. Reconstruction in AAFD patients who present with failure of the deltoid ligament and talar tilt requires rebalancing of the foot and tibiotalar joints. Surgical management typically depends on the presence of tibiotalar arthritis. In the absence of tibiotalar arthritis, reconstruction of the deltoid with a peroneus longus autograft is an option for these

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FIGURE 43.4 Patient with a hindfoot valgus deformity of the left heel on a hindfoot alignment view radiograph (A). Correction of the same hindfoot valgus deformity using a medializing calcaneal osteotomy with two cannulated screws (B). The weightbearing axis of the leg is now centered over the calcaneus. The goal of the surgeon should be to create a clinically straight heel with the hindfoot moment arm in 0 5 mm of varus. Lateral weightbearing radiograph of the same flatfoot with collapse of the medial longitudinal arch (C). Correction of the medial longitudinal arch in this foot after a medializing calcaneal osteotomy, lateral column lengthening, and modified Lapidus procedure (D).

patients [43,44]. Otherwise, consideration should be given to performing a total ankle arthroplasty or tibiotalar arthrodesis in patients with significant arthritis. These procedures for the treatment of deltoid ligament failure can be done in conjunction with an MCO or LCL to correct the associated hindfoot valgus deformity.

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Forefoot external rotation

As the flatfoot deformity progresses, the navicular externally rotates around the talus resulting in the forefoot external rotation deformity as seen on an anteroposterior view radiograph. Clinically, patients with pes planovalgus present with the “too many toes” sign, which is a clinical test primarily used to assess the hindfoot valgus deformity but is exacerbated by forefoot external rotation. Standing from behind the patient, the examiner asks the patient to stand with his or her feet shoulder width apart. More toes can be seen lateral to the calf on the affected side than on the normal side. As with all flatfoot deformities, the examiner must determine whether the deformity is flexible or rigid since this will guide treatment.

43.3.1 Bony anatomy The forefoot external rotation deformity typically arises in the midfoot at the talonavicular joint. As the navicular externally rotates around the talar head, the articular surface of the talar head becomes more “uncovered” by the navicular as viewed from the anteroposterior perspective. Adequate understanding of the forefoot external rotation deformity requires weightbearing radiographs, which may differ substantially from the examiner’s clinical impression of the deformity. Simultaneously, the navicular generally subluxes dorsally with respect to the talus as well. This will be discussed in more detail below. A significant forefoot external rotation deformity has previously been defined as greater than 30% uncoverage of the talonavicular head on weightbearing anteroposterior radiographs (Fig. 43.5) [2,5]. This cutoff has been used to divide stage IIA AAFD, which include patients with less than 30% uncoverage of the talonavicular head, and stage IIB AAFD, which includes patients with greater than 30% talonavicular uncoverage [2,5]. The talonavicular uncoverage percent is the amount of medial articular surface arc length of the talar head that is not articulating with the navicular on a weightbearing AP radiograph divided by the total arc length of the articular surface of the talar head [12]. Other measurements that have been used to quantify the talonavicular external rotation deformity include the talonavicular coverage angle [45], the talar-first metatarsal angle [15], and more recently the incongruency angle [12]. The incongruency angle is the

FIGURE 43.5 Mild talonavicular external rotation deformity (,30% of the articular surface of the talus is not covered by the navicular) on a weightbearing AP radiograph in a patient with stage IIA AAFD (A). More severe external rotation deformity ( . 30% of the articular surface of the talus is not covered by the navicular), which occurs in stage IIB AAFD (B).

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angle that is created by the intersection of a line between the lateral extent of the articular surfaces of the navicular and talus and a line from the lateral aspect of the talar neck at its most narrow segment and the lateral extent of the talar articular surface [12]. These measurements of talonavicular external rotation have excellent intra- and interrater reliabilities [12]. Weightbearing CT scans may also be used to quantify the talonavicular external rotation deformity. No studies have been published on the use of weightbearing CT scans on forefoot external rotation in AAFD.

43.3.2 Ligament and tendon failure The bony talonavicular external rotation deformity evident on radiographs results from the failure of the superomedial calcaneonavicular ligament and posterior tibial tendon [25]. The superomedial calcaneonavicular ligament together with the inferior calcaneonavicular fibers comprise the spring ligament complex [46]. Pathology of the spring ligament complex has been shown to occur almost as frequently as posterior tibial tendinopathy in flatfoot pathology [25,46]. The superomedial fibers of the calcaneonavicular ligament run from the superomedial aspect of the sustentaculum tali and anterior facet of the calcaneus to the medial edge of the navicular [46]. This region of the spring ligament complex is composed of stronger and larger fibers than the inferior region [46,47]. Moderate to severe superomedial calcaneonavicular ligament attenuation on MRI was found in 74% of patients with AAFD [25]. Failure of the superomedial calcaneonavicular ligament allows the talar head to internally rotate resulting in external rotation of the talonavicular joint and forefoot [46]. Degeneration of the posterior tibial tendon also contributes to the external rotation deformity of the forefoot. The posterior tibial tendon originates in the deep compartment of the leg from the posterior fibula, tibia, and interosseous membrane and inserts on the navicular tuberosity in addition to its multiple insertions along the plantar aspects of the midfoot and metatarsal bases. Before the heel rise stage of stance, the posterior tibial muscle contracts, which reduces external rotation of the midfoot on the hindfoot and locks the transverse tarsal joint creating a rigid lever [48]. Insufficiency of the posterior tibial tendon, as occurs with significant degeneration or frank tearing (Fig. 43.6), results in loss of internal rotation at the talonavicular joint and contributes to forefoot external rotation.

43.3.3 Surgical reconstruction Indications for surgical correction of the talonavicular external rotation deformity are controversial but typically include greater than 30% uncoverage of the talar head or incongruency at the lateral talonavicular joint as measured on weightbearing AP radiographs [2,33,45,49]. The Evans LCL procedure is most commonly used to correct the forefoot external rotation deformity [2,10,50]. Calcaneocuboid distraction arthrodesis has also been used for this purpose although it has higher rates of nonunion and postoperative pain than the LCL [51]. Additionally, less than 50% of talonavicular motion and less than 40% eversion of the talonavicular joint were preserved following a calcaneocuboid distraction arthrodesis [52]. It is important to note that the lateral column does not progressively shorten in patients with AAFD and the perceived difference in the lateral column length is due to deformity at the talonavicular joint [53]. In addition to internally rotating the foot at the talonavicular joint, the LCL also plantarflexes the midfoot and partially corrects hindfoot valgus

FIGURE 43.6 Axial slices of a T1-weighted MRI in patient with posterior tibial tendon insufficiency superior to the tibial plafond (A) and at the level of the sustentaculum tali (B). The posterior tibial tendon has abnormal high signal intensity in the images (white arrows). Axial slice of a T1weighted MRI in patient with a normal posterior tibial tendon for comparison (white arrow) (C).

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FIGURE 43.7 Weightbearing AP (A) and lateral (B) radiographs, respectively, of a foot following an Evans-type lateral column lengthening (LCL). Weightbearing AP (C) and lateral (D) radiographs of a step-cut “Z” LCL. Both types of LCLs provide correction of the external rotation deformity at the talonavicular joint and may partially correct the hindfoot valgus deformity.

resulting in restoration of the medial longitudinal arch [10,34,42,54]. However, the LCL procedure is not without risks, and patients undergoing an LCL have a higher incidence of nonunion and radiographic progression of adjacent joint arthritis than patients undergoing a MCO alone [28]. Nonunion rates following Evans-type LCL are approximately 10% [55]. Lateral plantar foot pain in patients treated with an LCL is also common with the incidence being reported from 8% to 45% [5,51,56]. Another option for LCL is through a step-cut “Z” osteotomy as described by Vander Griend (Fig. 43.7) [9,57]. In a cadaveric flatfoot deformity model, a step-cut LCL osteotomy was shown to correct all of the talonavicular external rotation deformity and 60% of the hindfoot valgus deformity [42]. The Evans LCL is traditionally performed with an opening wedge osteotomy in the anterior calcaneus approximately 12 mm proximal to the calcaneocuboid joint [9,58]. An opening wedge graft approximately 4 8 mm at its

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widest point is used to correct the external rotation deformity [58]. Autograft and allograft wedges have similar rates of nonunion and loss of correction [55]. More recent studies have demonstrated that LCLs using porous titanium wedges also have acceptable radiographic and clinical outcomes at a mean followup of over 1 year [59,60]. The step-cut osteotomy is made with distal and proximal vertical cuts in the anterior calcaneus that are connected via a horizontal limb [57]. Studies have demonstrated similar improvements in clinical outcome scores between patients treated with an Evans and step-cut osteotomy [61,62]. Patients undergoing a step-cut osteotomy were found to have less time to union on CT scan and fewer nonunions compared with Evans osteotomy patients [61]. Because the step-cut osteotomy is a newer technique and technically more demanding, further work needs to be done comparing the two procedures. Following an Evans-type LCL procedure, plantar pressures in the lateral forefoot of a cadaveric model increased significantly, which may account for the high rates of lateral foot pain following the procedure [35]. In a study of patients who had undergone an LCL, 10 patients with lateral foot pain were compared with a control group of 10 patients without pain, and the group of patients who had lateral foot pain were found to have increased plantar pressures in the lateral midfoot [63]. One cadaveric study demonstrated that increments as small as 2 mm in the size of the LCL led to both significant correction of the midfoot external rotation deformity and a linear increase in lateral forefoot plantar pressures [64]. Another study found that each 1 mm increase in the size of the LCL led to a 7-degree change in the lateral incongruency angle and has the greatest contribution to correction of the talonavicular external rotation deformity during a flatfoot reconstruction [49]. Thus, some authors have advocated for using trial wedges to minimize overcorrection of the lateral column [65]. Fifth metatarsal stress fractures have also been reported after LCLs [66]. Special care must be taken to avoid overcorrection of the talonavicular external rotation deformity. Avoiding overcorrection of the external rotation deformity when utilizing an LCL as part of a flatfoot reconstruction is important to preserve hindfoot motion and prevent loss of eversion. Patient-reported outcomes are also significantly affected by the LCL procedure [67]. In a retrospective review of 55 patients, patients who were overcorrected into talonavicular internal rotation following an LCL for AAFD demonstrated significantly lower improvements in clinical outcomes than those who were left in slight external rotation [67]. An internally rotated position of the talonavicular joint likely exacerbates lateral column overload and postoperative pain [63,68]. Spring ligament reconstruction has also been proposed as a technique to correct external rotation of the talonavicular joint [46,69]. A biomechanical study comparing three spring ligament reconstructive procedures demonstrated that a combined superomedial and plantar technique was the most effective for restoring internal rotation at the talonavicular joint and inversion at the subtalar joint [46]. A later study utilizing a different technique for spring ligament reconstruction by the same authors demonstrated good correction in 13 patients who had persistent talonavicular external rotation (. 30 degrees) or talonavicular sag (. 10 degrees) despite utilization of both an MCO and an LCL [70]. This spring ligament reconstruction was performed via a navicular tunnel through which a peroneus longus autograft was passed and then secured through a calcaneal or tibial drill hole [70]. The authors reported significant improvement in clinical outcome scores and talonavicular external rotation following spring ligament reconstruction [70]. A recent biomechanical study demonstrated that securing the graft through a bone tunnel in the tibia rather than using a more anatomic reconstruction of the spring ligament complex with a bone tunnel in the calcaneus resulted in a more significant correction of the external rotation and eversion deformities associated with AAFD [71]. Finally, in patients with significant deformity at the talonavicular joint that cannot be corrected with these procedures or patients with inflexible deformity at the talonavicular joint, a talonavicular fusion may be necessary. However, fusion at the talonavicular joint results in decreased motion through the hindfoot and may accelerate arthrosis in adjacent joints [72].

43.4

Sag at the talonavicular joint

Sag at the talonavicular joint results in collapse of the medial longitudinal arch of the foot, which consists of the talus, navicular, cuneiforms, and first metatarsal. Examination of the foot demonstrates decreased height of the medial arch, which often touches the ground resulting in the development of calluses at this site. The examiner may be able to palpate the talar head at the plantar medial aspect of the foot and should assess whether the deformity is flexible. Patients with AAFD do not reconstitute their arch during heel-toe progression. The severity of the plantar sag at the talonavicular joint cannot be reliably assessed using clinical examination alone and is better quantified on weightbearing lateral radiographs or weightbearing CT of the foot.

43.4.1 Bony anatomy Lateral weightbearing radiographs are used to measure the plantar sag deformity at the talonavicular joint [48,73]. The lateral first tarsometatarsal (TMT) angle, or Meary’s angle, is an angle formed by the intersection of a line through the long

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FIGURE 43.8 Weightbearing CT scans demonstrating the numerous deformities associated with the adult-acquired flatfoot deformity, including (1) collapse of the medial longitudinal arch (A)—Meary’s angle is seen to be clearly apex plantar in this image; (2) hindfoot valgus deformity (B)—the calcaneus is in valgus compared with the position of the talus, which is centered under the tibial plafond; (3) talonavicular external rotation deformity (C). A 3D reconstruction of the foot showing the medial longitudinal arch as well as the hindfoot demonstrates how the various deformities contribute to the clinical picture of a patient with adult-acquired flatfoot deformity (D).

axis of the talus and a line through the long axis of the first metatarsal. In normal feet, the lateral talometarsal angle on weightbearing radiographs is between 4 degrees apex dorsal and 4 degrees apex plantar. In AAFD patients, deformity at the talonavicular is considered mild when the angle is between 4 15 degrees apex plantar, moderate when the angle is between 15 30 degrees apex plantar, and severe when the angle is greater than 30 degrees apex plantar [73]. Newer technologies such as weightbearing CT scans may be able to better isolate the severity of plantar sag at the talonavicular joint from confounding deformities such as hindfoot valgus and talonavicular external rotation (Fig. 43.8).

43.4.2 Ligamentous failure The inferior calcaneonavicular component of the spring ligament complex is the second most commonly injured structure in AAFD and is responsible for plantar sag at the talonavicular joint [4,25,70]. The inferior calcaneonavicular ligament is weaker and more narrow than the superomedial portion of the spring ligament complex and is an entirely fibrous structure [25]. It runs from the anterior sustentaculum tali to the middle of the plantar aspect of the navicular [46]. It acts as a sling on the plantar aspect of the talonavicular joint preventing plantarflexion of the talar head [25,46]. MRI has been used to evaluate the integrity of the spring ligament complex, and the two components of the complex can be assessed individually (Fig. 43.9) [25].

43.4.3 Surgical reconstruction Isolated reconstruction of the inferior calcaneonavicular ligament has not been reported. Early reconstructions used the superficial deltoid ligament to reconstruct the spring ligament complex, but there were concerns about

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FIGURE 43.9 Axial proton density weighted MR image demonstrating the inferior calcaneonavicular component of a normal spring ligament complex (thick white arrow) as well as the medioplantar oblique component of the spring ligament complex (thin white arrow) (A). An oblique coronal proton density MR image through the hindfoot showing the superomedial fibers of a normal spring ligament (thick white arrows) (B).

graft strength. More recent techniques have utilized the peroneus longus autograft for this purpose [46,70,74]. Choi et al. demonstrated that plantar reconstruction of the spring ligament did not provide as much correction (, 5 degrees) of the plantarflexion deformity at the talonavicular joint as an approach that reconstructed both the inferior and superomedial portions of the spring ligament [46]. More recent techniques utilize a peroneus longus autograft through bone tunnels in the navicular and tibia or calcaneus with good improvements in patient outcomes, the lateral talonavicular angle, and the talonavicular coverage angle [70]. A graft inserted into the calcaneus better mimics the anatomic path of the inferior calcaneonavicular ligament while a graft inserted into the tibia better follows the path of the superomedial calcaneonavicular ligament [70]. At this time, indications for the use of a spring ligament reconstruction are unclear and based on the surgeon’s experience. Techniques describing reconstruction of the spring ligament often cite the presence of persistent talonavicular external rotation (. 30 degrees) or talonavicular sag (. 10 degrees) after the use of an MCO and an LCL; however, there are no guidelines for preoperative planning [70]. In severe cases of talonavicular plantar sag or in rigid deformities, arthrodesis of the medial column may stabilize the medial longitudinal arch [2,48,75]. Previously, triple arthrodesis had been advocated for correction of severe deformity at the talonavicular joint [2]. Isolated fusion of the talonavicular joint has been shown to correct significant deformity [7,76]. Midfoot arthrodesis or combined subtalar and naviculocuneiform fusions may also be used to correct talonavicular sag although many surgeons reserve these procedures for rigid flatfoot deformities [75,77]. Such fusions decrease hindfoot motion and likely overall function.

43.5

Failure of the posterior tibial tendon

The posterior tibial tendon acts as an inverter and plantarflexor of the foot. During the stance phase of the gait cycle, the calcaneus moves into valgus and the cuboid externally rotates the forefoot leading to collapse of the longitudinal arch of the foot [78]. As the medial longitudinal arch collapses, the posterior tibial tendon eccentrically contracts to stabilize the arch and control eversion of the foot [78]. Subsequently, the hindfoot inverts during toe-off and the posterior tibial tendon pulls the foot into a slight varus position allowing the transverse tarsal joint to lock and act as a rigid lever arm for push-off [78]. Biomechanical studies have demonstrated that the posterior tibial tendon provides significant stability to the arch [24,79]. When the posterior tibial tendon is not loaded, arch height significantly decreases and external rotation of the forefoot significantly increases [79]. In AAFD, failure of the posterior tibial tendon is a degenerative process resulting in dysfunction of the tendon [2]. Tendon pathology is typically seen at the level of the medial malleolus or just distal to it, which may be related

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to altered biomechanics caused by preexisting bony abnormalities, ligamentous laxity, obesity, and/or the relative hypovascularity of the tendon in this region [2,80]. The hypovascularity of this region may explain why patients with hypertension and diabetes mellitus are more susceptible to developing AAFD [80].

43.5.1 Clinical assessment To assess the integrity of the posterior tibial tendon, the single heel rise test can be performed. The patient stands on one foot with a straight knee while balancing against the wall. The contralateral foot is raised off the ground. The patient is then asked to lift their heel off the ground and go onto their toes. A positive test is one in which the patient is unable to lift the heel off the ground or if normal heel inversion is absent [2]. Some patients with posterior tibial tendon insufficiency may be to lift their heel off the ground without the heel moving into varus as typically occurs during a heel rise [2]. Arthritis, Achilles tendon ruptures, and previous fusions may give false-positive results [2]. Strength testing of the posterior tibial tendon with plantarflexion and inversion may not reveal the extent of the deficit due to the contribution of the tibialis anterior tendon to inversion [48]. In addition to the single heel rise, the posterior tibial tendon edema sign is a useful test for dysfunction of the tendon [81]. The posterior tibial tendon edema sign occurs when patients have pitting edema along the length of the posterior tibial tendon sheath that is not due to acute trauma or associated with edema in other parts of the foot [81]. This physical examination finding was highly sensitive (86%) and specific (100%) when compared with MRI findings [81]. Patients may also complain of tenderness upon palpation of the posterior tibial tendon from the medial malleolus to the navicular tuberosity, and in some cases, more proximally at the musculotendinous junction [48].

43.5.2 MRI assessment MRI is a useful imaging modality in the evaluation of the posterior tibial tendon [25,82 84]. In a study comparing patients with AAFD to patients with normal feet, over 90% of patients with AAFD have grade II or greater tears in their posterior tibial tendon whereas none of the patients with normal feet had grade II tears in their tendon [25]. In patients with AAFD, atrophy of the posterior tibial tendon and compensatory FDL hypertrophy have been confirmed on MRI assessments [84].

43.5.3 Surgical reconstruction Insufficiency of the posterior tibial tendon is typically treated with a FDL tendon transfer although the flexor hallucis longus tendon has been utilized as well [29 31,36,37,39,85]. The FDL transfer has long been used to augment a deficient posterior tibial tendon as the FDL tendon courses near and fires in phase with the posterior tibial tendon [86,87]. However, the PTT is approximately 3.5 times stronger than the FDL, and the FDL tendon after transfer must also counter the force of the peroneus brevis, which is the antagonist of the PTT and 1.5 times the strength of the FDL [88]. The FDL tendon transfer is typically combined with an MCO to protect the tendon transfer [10]. The FDL transfer is performed through a medial incision directly over the posterior tibial and FDL tendons. The FDL tendon is found interior to the posterior tibial tendon sheath and freed distally to the knot of Henry. The proximal FDL tendon is then passed through the navicular from plantar to dorsal and, at the end of the case to ensure correct tensioning of the FDL tendon, sutured to the stump of the posterior tibial tendon with the ankle in neutral and the foot slightly inverted. Special consideration should be given to younger patients presenting with a flatfoot deformity as the posterior tibial tendon is often relatively healthy in this population; consequently, an FDL tendon transfer may not be necessary [89].

43.6

Gastrocnemius and Achilles tightness

Contraction of the gastrocnemius-soleus complex or Achilles tendon is common in AAFD. Increased pull from the Achilles tendon, for example, due to contraction of the tendon, results in increased force across the tibiotalar joint [78]. Because the tibiotalar joint lies between the calcaneal tuberosity and forefoot, which are two weightbearing portions of the foot, the force of the tibiotalar joint is transmitted to the forefoot via the medial longitudinal arch [78]. As the foot moves from the heel strike phase of the gait cycle to the toe-off phase, the gastrocnemius-soleus complex eccentrically contracts to control dorsiflexion of the tibia. Subsequently, the knee extends, and the gastrocnemius, through the Achilles tendon, exerts increased force on the tibiotalar joint [78]. In the setting of a compromised posterior tibial

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FIGURE 43.10 Clinical photograph assessing ankle dorsiflexion with the knee straight (A) or with the knee flexed (B). Improved ankle dorsiflexion with the knee flexed suggested gastrocnemius tightness. If ankle dorsiflexion does not improve with flexion of the knee, then the patient has Achilles tendon tightness, which cannot be corrected with gastrocnemius lengthening alone.

tendon and medial longitudinal arch, the foot does not form a rigid lever arm for toe-off leading to the transmission of force to the midfoot rather than the forefoot, increased stress and collapse of the medial longitudinal arch, and an inefficient gait [90]. Thus, contraction of the gastrocnemius-soleus complex or Achilles tendon causes further flattening of the medial longitudinal arch and, following reconstruction, places excessive stress on the operative foot.

43.6.1 Clinical assessment Because tightness in the gastrocnemius muscle or Achilles tendon is common in AAFD, physical examination of the flatfoot patient should include assessment of these structures. The gastrocnemius originates on the medial and lateral condyles of the posterior femur and crosses the knee and ankle joints. Thus, to evaluate for isolated gastrocnemius muscle vs Achilles tendon tightness, the Silfverskiold test is performed, in which ankle dorsiflexion is tested with the knee extended and the knee at 90 degrees of flexion. If ankle dorsiflexion does not improve with knee flexion, then the patient has Achilles tightness that cannot be corrected with a gastrocnemius lengthening alone (Fig. 43.10) [2].

43.6.2 Surgical treatment A gastrocnemius recession or tendo-Achilles lengthening is typically performed as part of a flatfoot reconstruction as tightness of these soft tissues is common among patients with AAFD and can put excess pressure on the repair if not addressed [7]. In cases of isolated gastrocnemius tightness, a gastrocnemius recession is done at the beginning of the case to gain appropriate ankle dorsiflexion [58]. A triple-cut tendo-Achilles lengthening procedure is performed when the Silfverskiold test reveals that tightness is from a contracted Achilles tendon [2,58] or a gastrocnemius recession is not sufficient intraoperatively to improve ankle dorsiflexion.

43.7

Medial arch eversion

Eversion of the medial longitudinal arch is a secondary deformity that results from persistent hindfoot valgus [91]. This deformity may present as elevation of the first ray or forefoot varus at the transverse tarsal joint [91]. Frequently, these patients present with hypermobility of the first ray [92].

43.7.1 Clinical assessment Evaluation of the medial arch eversion deformity should occur with the patient in the seated position and the hindfoot placed in a neutral position with the navicular centered on the talar head [91]. The examiner palpates the medial border

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of the foot and the relationship of the navicular tuberosity to the talar head [91]. As the ankle is brought into neutral dorsiflexion through pressure on the fourth and fifth metatarsal heads with the hindfoot in neutral, elevation of the first metatarsal with reference to the fifth metatarsal determines the degree of forefoot varus [91]. Additionally, stability of the first ray is tested by dorsiflexing and plantarflexing the first metatarsal while the other four metatarsals are stabilized [91]. Crepitus, pain, or excessive motion suggests an unstable first ray [91].

43.7.2 Bony anatomy The bony anatomy of the first TMT joint is assessed on weightbearing radiographs or CT scans. Lateral radiographs may reveal plantar-gapping or subluxation at the first TMT joint, which indicates instability of this joint [91]. Additionally, degenerative changes at the first TMT joint such as joint space narrowing, subchondral sclerosis, or osteophytes may be evident [91]. Weightbearing CT scans may be helpful in evaluating the location of the medial arch eversion deformity and better assessing degenerative changes at the first TMT joint.

43.7.3 Ligamentous failure No specific ligamentous failures have been identified that lead to the midfoot collapse associated with the medial arch eversion deformity. This deformity results from persistent hindfoot valgus, which places significant strain on the medial arch leading to progressive degeneration of the soft tissue structures that support the midfoot. In addition to the bony collapse associated with this deformity, the plantar ligaments that span the medial cuneiform and first ray as well as its joint capsule become insufficient [4]. Specific joint capsules of the talonavicular, naviculocuneiform, and/or TMT joints maybe compromised as well.

43.7.4 Surgical reconstruction Reconstruction of the first TMT joint is performed as part of a flatfoot reconstruction after correction of the hindfoot deformity if there is residual eversion of the midfoot. This deformity should be assessed preoperatively both clinically and radiographically in addition to using intraoperative simulated weightbearing lateral radiographs [91]. Residual eversion can be addressed through a first TMT fusion or Cotton osteotomy. A modified Lapidus technique using two crossing cortical screws is indicated when there is significant plantar-gapping at the first TMT joint, a concomitant hallux valgus deformity, or first TMT joint instability or arthritis [93 95]. With newer techniques, nonunion rates following first TMT fusion are low and have been reported to be less than 3% [93]. If residual medial arch eversion persists after correction of the hindfoot deformity and the patient does not meet these criteria for a modified Lapidus procedure, then a Cotton osteotomy may be performed, which is a joint-sparing procedure [91]. The Cotton osteotomy is a plantarflexion opening wedge osteotomy in the medial cuneiform that helps to plantarflex the first ray and restore the tripod of the foot [91]. The Cotton osteotomy is contraindicated if there is significant subluxation or arthritis of the first TMT joint or if the deformity requires greater than an 8 10 mm wedge [91]. One retrospective study demonstrated significant improvement in radiographic parameters following a Cotton osteotomy during a flatfoot reconstruction [96]. There was no evidence of radiographic instability of the first TMT joint following the procedure [96]. In cases of plantar gapping at the naviculocuneiform joint, a naviculocuneiform fusion is more appropriate as the modified Lapidus procedure and Cotton osteotomy do not address this deformity. A study looking at the results of ten patients with stage III AAFD who underwent reconstruction utilizing a subtalar and naviculocuneiform fusion demonstrated radiographic correction and patient satisfaction at a minimum of one year postoperatively [75]. This procedure is likely best reserved for rigid deformities or if there is specific deformity at the naviculocuneiform joint. Difficulties with this procedure include the ability to address the articulation of the navicular with all three cuneiforms and the potential for nonunion.

43.8

Special considerations

Recent papers have demonstrated that flatfoot reconstructions can be successfully performed in high-risk patient groups [3,89,97]. Adolescents with flatfoot deformity have been successfully treated with a combination of an MCO, LCL, and gastrocnemius recession [89]. In a study of sixteen adolescent and young adult patients with idiopathic flatfoot who underwent flatfoot reconstructions, all but one were able to return to sports and physical activity [89]. This population, however, requires a FDL tendon transfer less often than older patients as their posterior tibial tendon is frequently intact and has not undergone significant degeneration [89].

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Patients who are overweight (BMI greater than or equal to 25 but less than 30) or obese (BMI greater than or equal to 30) may be at higher risk of having their flatfoot reconstruction fail due to increased load across the medial longitudinal arch. However, short-term clinical outcomes following flatfoot reconstructions in these patients do not show results that are different from normal-weight patients [3]. Similarly, older patients (age greater than or equal to 65 years old) do not show significantly different postoperative clinical outcome scores at 2 years postoperatively or have an increased number of subsequent surgical procedures following flatfoot reconstructions than younger patients [97]. Patients in the older group did not require significantly more conversions to fusion procedures than younger patients after reconstruction of their flatfeet [97]. Consequently, joint-preserving reconstruction of the adult-acquired flatfoot in higher risk populations remains a viable option and an alternative to fusion in these patients [3,97].

43.9

Future biomechanical studies and conclusion

Future work in the biomechanics of flat foot reconstruction should focus on correlating specific ligamentous and soft tissue failures with their associated deformities. This will help clarify indications for specific procedures such as the spring ligament reconstruction. The utilization of weightbearing CT scans is effective to better assess osseous abnormalities in AAFD and will assist in our understanding of the bony deformities that result from improper loading of the foot [98]. Additional work using patient-reported outcomes following reconstruction of the flatfoot are necessary to improve surgical techniques and specify indications for joint-preserving procedures. In general, the flatfoot deformity can be summarized as peritalar subluxation with the ligamentous, tendinous, and bony structures in close proximity to the talus failing in ways that lead to progressive deformity. This chapter has reviewed how insufficiencies in those structures result in hindfoot valgus, forefoot external rotation, sag at the talonavicular joint, and medial arch eversion in patients with AAFD. Previous biomechanical studies have attempted to correlate specific anatomic structures that, when insufficient, lead to these abnormalities in the foot. Failure of the talocalcaneal interosseous ligament, superomedial calcaneonavicular ligament, inferior calcaneonavicular ligament, posterior tibial tendon, and deltoid ligament as well as instability at the first TMT joint and gastrocnemius muscle and Achilles tendon tightness contribute to the altered mechanics of the foot in AAFD. Reconstruction of the adult-acquired flatfoot must be individualized to the patient. Understanding how the specific deformities affect the biomechanics of the foot allows the surgeon to tailor the reconstruction to each patient—reestablishing proper alignment, ideally a flexible foot, and a more normal distribution of pressures during the gait cycle. Ultimately, the goal should be to decrease pain and improve function in patients who suffer from AAFD regardless of the severity of the deformities.

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Foot Ankle Int 1996;17:95 102. Available from: https://doi.org/10.1177/107110079601700207. [48] Pinney SJ, Lin SS. Current concept review: acquired adult flatfoot deformity. Foot Ankle Int 2006;27:66 75. [49] Chan JY, Greenfield ST, Soukup DS, Do HT, Deland JT, Ellis SJ. Contribution of lateral column lengthening to correction of forefoot abduction in stage IIb adult acquired flatfoot deformity reconstruction. Foot Ankle Int 2015;36:1400 11. Available from: https://doi.org/10.1177/ 1071100715596607. [50] Moseir-LaClair S, Pomeroy G, Manoli 2nd A. Intermediate follow-up on the double osteotomy and tendon transfer procedure for stage II posterior tibial tendon insufficiency. Foot Ankle Int 2001;22:283 91. [51] Thomas RL, Wells BC, Garrison RL, Prada SA. Preliminary results comparing two methods of lateral column lengthening. Foot Ankle Int 2001;22:107 19. Available from: https://doi.org/10.1177/107110070102200205. [52] Deland JT, Otis JC, Lee K-T, Kenneally SM. Lateral column lengthening with calcaneocuboid fusion: range of motion in the triple joint complex. Foot Ankle Int 1995;16:729 33. Available from: https://doi.org/10.1177/107110079501601111. [53] Basta NW, Mital MA, Bonadio O, Johnson A, Kang SY, O’Connor J. A comparative study of the role of shoes, arch supports, and navicular cookies in the management of symptomatic mobile flat feet in children. Int Orthop 1977;1:143 8. Available from: https://doi.org/10.1007/ BF00576318. [54] Dumontier TA, Falicov A, Mosca V, Sangeorzan B. Calcaneal lengthening: investigation of deformity correction in a cadaver flatfoot model. Foot Ankle Int 2005;26:166 70. [55] Vosseller JT, Ellis SJ, O’Malley MJ, Elliott AJ, Levine DS, Deland JT, et al. Autograft and allograft unite similarly in lateral column lengthening for adult acquired flatfoot deformity. HSS J 2013;9:6 11. Available from: https://doi.org/10.1007/s11420-012-9317-5. [56] van der Krans A, Louwerens JWK, Anderson P. Adult acquired flexible flatfoot, treated by calcaneocuboid distraction arthrodesis, posterior tibial tendon augmentation, and percutaneous Achilles tendon lengthening: a prospective outcome study of 20 patients. Acta Orthop 2006;77:156 63. Available from: https://doi.org/10.1080/17453670610045858. [57] Vander Griend R. Lateral column lengthening using a “Z” osteotomy of the calcaneus. Tech Foot Ankle Surg 2008;7:257 63. Available from: https://doi.org/10.1097/BTF.0b013e318183a0df. [58] Conti MS, Deland JT, Ellis SJ. Stage IIB flatfoot reconstruction using literature-based equations for heel slide and lateral column lengthening. Tech Foot Ankle Surg 2017;16:153 66. Available from: https://doi.org/10.1097/BTF.0000000000000164. [59] Moore SH, Carstensen SE, Burrus MT, Cooper T, Park JS, Perumal V. Porous titanium wedges in lateral column lengthening for adult-acquired flatfoot deformity. Foot Ankle Spec 2017;. Available from: https://doi.org/10.1177/1938640017735890 1938640017735890. [60] Gross CE, Huh J, Gray J, Demetracopoulos C, Nunley JA. Radiographic outcomes following lateral column lengthening with a porous titanium wedge. Foot Ankle Int 2015;36:953 60. Available from: https://doi.org/10.1177/1071100715577788. [61] Saunders SM, Ellis SJ, Demetracopoulos CA, Marinescu A, Burkett J, Deland JT. Comparative outcomes between step-cut lengthening calcaneal osteotomy vs traditional evans osteotomy for stage IIB adult-acquired flatfoot deformity. Foot Ankle Int 2018;39:18 27. Available from: https://doi.org/10.1177/1071100717732723. [62] Demetracopoulos CA, Nair P, Malzberg A, Deland JT. Outcomes of a stepcut lengthening calcaneal osteotomy for adult-acquired flatfoot deformity. Foot Ankle Int 2015;36:749 55. Available from: https://doi.org/10.1177/1071100715574933. [63] Ellis SJ, Yu JC, Johnson AH, Elliott A, OʼMalley M, Deland J. Plantar pressures in patients with and without lateral foot pain after lateral column lengthening. J Bone Jt Surg Am 2010;92:81 91. Available from: https://doi.org/10.2106/JBJS.H.01057. [64] Oh I, Imhauser C, Choi D, Williams B, Ellis S, Deland J. Sensitivity of plantar pressure and talonavicular alignment to lateral column lengthening in flatfoot reconstruction. J Bone Jt Surg Am 2013;95:1094 100. [65] Ellis SJ, Williams BR, Garg R, Campbell G, Pavlov H, Deland JT. Incidence of plantar lateral foot pain before and after the use of trial metal wedges in lateral column lengthening. Foot Ankle Int 2011;32:665 73. Available from: https://doi.org/10.3113/FAI.2011.0665. [66] Davitt JS, Morgan JM. Stress fracture of the fifth metatarsal after Evans’ calcaneal osteotomy: a report of two cases. Foot Ankle Int 1998;19:710 12. [67] Conti MS, Chan JY, Do HT, Ellis SJ, Deland JT. Correlation of postoperative midfoot position with outcome following reconstruction of the stage II adult acquired flatfoot deformity. Foot Ankle Int 2015;. Available from: https://doi.org/10.1177/1071100714564217. [68] Davitt JS, MacWilliams BA, Armstrong PF. Plantar pressure and radiographic changes after distal calcaneal lengthening in children and adolescents. J Pediatr Orthop 2001;21:70 5. [69] Deland JT. The adult acquired flatfoot and spring ligament complex. Pathology and implications for treatment. Foot Ankle Clin 2001;6:129 35 vii. [70] Williams BR, Ellis SJ, Deyer TW, Pavlov H, Deland JT. Reconstruction of the spring ligament using a peroneus longus autograft tendon transfer. Foot Ankle Int 2010;31:567 77. Available from: https://doi.org/10.3113/FAI.2010.0567.

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[71] Baxter JR, LaMothe JM, Walls RJ, Prado MP, Gilbert SL, Deland JT. Reconstruction of the medial talonavicular joint in simulated flatfoot deformity. Foot Ankle Int 2015;36:424 9. Available from: https://doi.org/10.1177/1071100714558512. [72] Suckel A, Muller O, Herberts T, Langenstein P, Reize P, Wulker N. Talonavicular arthrodesis or triple arthrodesis: peak pressure in the adjacent joints measured in 8 cadaver specimens. Acta Orthop 2007;78:592 7. Available from: https://doi.org/10.1080/17453670710014275. [73] Pedowitz, Kovatis. Flatfoot in the adult. J Am Acad Orthop Surg 1995;3:293 302. [74] Deland JT, Arnoczky SP, Thompson FM. Adult acquired flatfoot deformity at the talonavicular joint: reconstruction of the spring ligament in an in vitro model. Foot Ankle 1992;13:327 32. [75] Barg A, Brunner S, Zwicky L, Hintermann B. Subtalar and naviculocuneiform fusion for extended breakdown of the medial arch. Foot Ankle Clin 2011;16:69 81. Available from: https://doi.org/10.1016/j.fcl.2010.11.004. [76] O’Malley MJ, Deland JT, Lee KT. Selective hindfoot arthrodesis for the treatment of adult acquired flatfoot deformity: an in vitro study. Foot Ankle Int 1995;16:411 17. Available from: https://doi.org/10.1177/107110079501600706. [77] Greisberg J, Assal M, Hansen ST, Sangeorzan BJ. Isolated medial column stabilization improves alignment in adult-acquired flatfoot. Clin Orthop Relat Res 2005;197 202. [78] Van Boerum DH, Sangeorzan BJ. Biomechanics and pathophysiology of flat foot. Foot Ankle Clin 2003;8:419 30. [79] Kitaoka HB, Luo ZP, An KN. Effect of the posterior tibial tendon on the arch of the foot during simulated weightbearing: biomechanical analysis. Foot Ankle Int 1997;18:43 6. Available from: https://doi.org/10.1177/107110079701800109. [80] Holmes GB, Mann RA. Possible epidemiological factors associated with rupture of the posterior tibial tendon. Foot Ankle 1992;13:70 9. [81] DeOrio JK, Shapiro SA, McNeil RB, Stansel J. Validity of the posterior tibial edema sign in posterior tibial tendon dysfunction. Foot Ankle Int 2011;32:189 92. Available from: https://doi.org/10.3113/FAI.2011.0189. [82] Feighan J, Towers J, Conti S. The use of magnetic resonance imaging in posterior tibial tendon dysfunction. Clin Orthop Relat Res 1999;23 38. [83] Conti S, Michelson J, Jahss M. Clinical significance of magnetic resonance imaging in preoperative planning for reconstruction of posterior tibial tendon ruptures. Foot Ankle 1992;13:208 14. [84] Wacker J, Calder JDF, Engstrom CM, Saxby TS. MR morphometry of posterior tibialis muscle in adult acquired flat foot. Foot Ankle Int 2003;24:354 7. Available from: https://doi.org/10.1177/107110070302400409. [85] Sammarco GJ, Hockenbury RT. Treatment of stage II posterior tibial tendon dysfunction with flexor hallucis longus transfer and medial displacement calcaneal osteotomy. Foot Ankle Int 2001;22:305 12. [86] Jahss MH. Spontaneous rupture of the tibialis posterior tendon: clinical findings, tenographic studies, and a new technique of repair. Foot Ankle 1982;3:158 66. [87] Johnson KA. Tibialis posterior tendon rupture. Clin Orthop Relat Res 1983;140 7. [88] Silver RL, de la Garza J, Rang M. The myth of muscle balance. A study of relative strengths and excursions of normal muscles about the foot and ankle. J Bone Jt Surg Br 1985;67:432 7. [89] Oh I, Williams BR, Ellis SJ, Kwon DJ, Deland JT. Reconstruction of the symptomatic idiopathic flatfoot in adolescents and young adults. Foot Ankle Int 2011;32:225 32. Available from: https://doi.org/10.3113/FAI.2011.0225. [90] Mosier SM, Pomeroy G, Manoli A. Pathoanatomy and etiology of posterior tibial tendon dysfunction. Clin Orthop Relat Res 1999;12 22. [91] McCormick JJ, Johnson JE. Medial column procedures in the correction of adult acquired flatfoot deformity. Foot Ankle Clin 2012;17:283 98. Available from: https://doi.org/10.1016/j.fcl.2012.03.003. [92] Cowie S, Parsons S, Scammell B, McKenzie J. Hypermobility of the first ray in patients with planovalgus feet and tarsometatarsal osteoarthritis. Foot Ankle Surg 2012;18:237 40. Available from: https://doi.org/10.1016/j.fas.2012.01.004. [93] Mani SB, Lloyd EW, MacMahon A, Roberts MM, Levine DS, Ellis SJ. Modified lapidus procedure with joint compression, meticulous surface preparation, and shear-strain-relieved bone graft yields low nonunion rate. HSS J 2015;11:243 8. Available from: https://doi.org/10.1007/ s11420-015-9462-8. [94] Sangeorzan BJ, Hansen ST. Modified lapidus procedure for hallux valgus. Foot Ankle Int 1989;9:262 6. Available from: https://doi.org/ 10.1177/107110078900900602. [95] Lapidus PW. A quarter of a century of experience with the operative correction of the metatarsus varus primus in hallux valgus. Bull Hosp Jt Dis 1956;17:404 21. [96] Aiyer A, Dall GF, Shub J, Myerson MS. Radiographic correction following reconstruction of adult acquired flat foot deformity using the cotton medial cuneiform osteotomy. Foot Ankle Int 2016;37:508 13. Available from: https://doi.org/10.1177/1071100715620894. [97] Conti MS, Jones MT, Savenkov O, Deland JT, Ellis SJ. Outcomes of reconstruction of the stage II adult-acquired flatfoot deformity in older patients. Foot Ankle Int 2018;. Available from: https://doi.org/10.1177/1071100718777459 107110071877745. [98] de Cesar Netto C, Schon LC, Thawait GK, da Fonseca LF, Chinanuvathana A, Zbijewski WB, et al. Flexible adult acquired flatfoot deformity: comparison between weight-bearing and non-weight-bearing measurements using cone-beam computed tomography. J Bone Jt Surg Am 2017;99:e98. Available from: https://doi.org/10.2106/JBJS.16.01366.

Chapter 44

Cavus Foot Reconstructions Nathan Kiewiet1 and William Braaksma2 1

Orthopedic Health of Kansas City, North Kansas City, MO, United States, 2Orthopedic Associates of Port Huron, Port Huron, MI, United States

Abstract Pes cavus describes a deformity of the foot resulting in a high arch that remains stiff with weight bearing. Cavus alignment can be caused by mechanical factors of the hindfoot, forefoot, or both. Cavus alignment leads to consistent biomechanical consequences regardless of the cause of the deformity. These mechanical changes lead to dysfunction and symptoms. Treatment for cavus foot alignment can range from conservative management with inserts or bracing to surgical intervention including cavus reconstructions. Cavus reconstructions are intended to realign the anatomy to create a more balanced foot with more normal mechanics, with the overall aim of reducing the patient’s symptoms.

44.1

Introduction

Pes cavus is a foot type characterized by a high arch and altered foot mechanics. Even during weight bearing, cavus feet maintain their high arched shape. Pes cavus alignment can be caused predominantly by deformities of the hindfoot, forefoot, or a combination of both. The etiology of cavus deformities are typically neurologic, traumatic, residual clubfoot, or idiopathic in nature. Owing to varying components of deformity in the hindfoot and forefoot, a precise radiographic definition of pes cavus is difficult to ascertain. Cavus alignment must be viewed on a spectrum with the underlying abnormality being that of an elevated longitudinal arch.

44.2

Etiology

44.2.1 Traumatic Pes cavus can be the result of a previous traumatic injury. Deep posterior compartment syndrome may cause contracture and overpull of the tibialis posterior and the flexor digitorum longus muscles creating a cavovarus and equinus position. Talar neck malunion may cause a fixed varus position of the talonavicular, subtalar, and calcaneocuboid joints [1]. Traumatic injury causing loss of function of the peroneal nerve can result in peroneal muscle weakness. Unopposed action of the tibialis posterior and flexor digitorum longus due to peroneal weakness causes hindfoot and forefoot varus.

44.2.2 Neurologic Neurologic dysfunction can lead to muscle imbalances in agonist antagonist pairs. Brewerton et al. found subtle neurologic effects in 66% of patients in a series on the cause of pes cavus [2]. Cavus foot deformities caused by the hereditary motor and sensory neuropathies (HMSN) are mostly motor-driven [3]. Charcot-Marie-Tooth (CMT) disease is a group of disorders caused by defects in any of several constituent proteins of the myelin sheath of a peripheral nerve. It is a subgroup of the HMSNs that results in relatively strong tibialis posterior and peroneus longus muscles with weak peroneus brevis and tibialis anterior muscles. This muscle imbalance results in hindfoot varus and forefoot valgus. CMT patients have hypertrophy of the peroneus longus muscle creating an imbalance

Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00003-2 © 2023 Elsevier Inc. All rights reserved.

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TABLE 44.1 Etiology of pes cavus. Traumatic Fracture malunion Peroneal nerve injury Compartment syndrome Neurologic Hereditary motor and sensory neuropathies (e.g., Charcot-Marie-Tooth disease) Cerebral palsy Stroke Anterior horn cell disease Spinal cord lesions Residual clubfoot Idiopathic

with the tibialis anterior muscle [4]. CMT represents the single most common diagnosis associated with pes cavus alignment. It is the most common inheritable defect of peripheral nerve, however its unusual patterns of weakness continues to be poorly understood [5]. The etiology of pes cavus encompasses neurologic, traumatic, deformity, and idiopathic considerations (Table 44.1). Neurologic abnormalities with more proximal ramifications can also cause cavus foot alignment. Cavus deformities that are unilateral or show rapid progression require an aggressive work up to rule out a correctable lesion of the spinal cord. Clinical findings may also include hyperreflexia, clonus, or significant asymmetry in motor pattern.

44.2.3 Residual clubfoot Cavus alignment is one of the components of congenital clubfoot deformity. Untreated or partially treated clubfeet can result in persistent adult cavovarus and equinus deformity. This persistent deformity will result in fixed hindfoot varus and forefoot valgus. Residual clubfoot is typically a result of failure of early casting to correct the first ray as described by Ponseti [6].

44.2.4 Idiopathic Even though there are several known causes of pes cavus, there are numerous reported cases of cavus alignment occurring without an identifiable underlying cause, and these are usually symmetric. Most of these cases present in a more subtle form which is likely genetic and not neurologic in nature, and commonly presents in adults [7].

44.3

Clinical presentation and associated pathology

Patients with pes cavus deformity present with a variety of complaints and clinical findings. Owing to the structure of the foot, pressure is placed through the plantar aspect of the first metatarsal head, the lateral border of the foot, and the heel [8]. These biomechanical changes will cause pain at a variety of locations for the patient. Patients may present with pain at the plantar aspect of the first metatarsal head, the lesser metatarsal heads, or on the lateral foot at the base of the fifth metatarsal and cuboid [8]. Increased stress through the lateral border of the foot can lead to stress fractures of the fifth metatarsal and, less commonly, of the other lesser metatarsals. Saxena et al. has shown a correlation between cavus foot alignment and stress fractures at the base of the fourth metatarsal [9]. Hindfoot varus can result in overload of the soft tissues of the lateral ankle, including the lateral ankle ligaments and the peroneal tendons. Recurrent ankle instability can be caused by varus hindfoot alignment combined with the medially displaced mechanical pull of the Achilles tendon, leading to overload of the lateral ankle ligaments and instability [10]. Varus hindfoot alignment also creates chronic stress on the peroneal tendons which can lead to recurrent instability or dislocation of the tendons. Other chronic peroneal pathology can include tendonitis, splitting or tearing, and painful os peroneum syndrome with fragmentation and pain [11]. With time, chronic degenerative changes of the

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tibiotalar joint can develop due to varus hindfoot alignment. These changes can result in varus tilt of the talus within the tibiotalar joint [10]. Painful metatarsalgia can develop due to distal displacement of the plantar fat pad as a result of cavus foot deformity and extension of the metatarsophalangeal (MTP) joints combined with equinus contracture [10]. Painful calluses can develop beneath the first metatarsal head, the fifth metatarsal head, or the base of the fifth metatarsal. Chronic equinus contracture and contracture through the plantar fascia can cause painful plantar fasciitis. Chronic varus stress and external rotation at the ankle can lead to increased varus stress at the knee. The patient may present with varus strain at the knee, increased lateral collateral knee ligament strains, and iliotibial band syndrome. With chronic varus strain of the knee, medial compartment arthritis may develop with time. Stress related disorders such as stress fractures of the medial malleolus, tibia, or fibula may also develop due to the decrease in shock absorption produced with cavus deformity.

44.4

Physical exam

The physical exam starts with observation of the patient standing and their gait. Specific attention should be paid to ground contact of the foot, the position of the heel, and the position of the toes. Standing alignment of the cavus foot will reveal a “peek-a-boo” heel sign when observed from the front due to hindfoot varus alignment. Observation from behind the patient will also reveal hindfoot varus alignment. Coleman block testing can be used to assess the ability of the hindfoot to correct to the appropriate valgus position (Fig. 44.1). Coleman block testing to assess the rigidity of the hindfoot deformity is invaluable. With a forefoot-driven cavus deformity, Coleman block testing will result in notable correction of the hindfoot from varus to valgus consistent with a flexible hindfoot. With a hindfoot-driven cavus deformity, Coleman block testing will not result in correction of the hindfoot varus consistent with a rigid hindfoot deformity. With gait, any worsening of the hindfoot varus alignment during weight transfer should be noted. Careful attention should be paid for any signs of drop foot and toe deformity during gait. With seated examination, passive and active range of motion can be tested. Motion at the ankle, subtalar, transverse tarsal, and MTP joints should be observed. Strength is assessed with dorsiflexion (or extension), plantarflexion (or flexion), inversion, and eversion. Specific testing of the ability of the peroneus longus to flex the first ray is important as this can help determine the potential for tendon transfer to assist in treatment. It is important to assess the rigidity of the cavus deformity as this may dictate treatment options. Attention should also be made to any calluses. Calluses are common at the first and fifth metatarsal heads, as well as along the lateral foot. Lesser toe deformities are assessed for calluses, wounds, and flexibility when considering treatment options (Refer to Fig. 44.1A for claw toe deformities). The physical exam is perhaps the most important aspect in determining a successful treatment plan. The surgeon must identify muscle imbalances and the flexibility of the deformity. As with any elective lower extremity procedure, an assessment of vascular status and peripheral pulses should be performed. If there is any concern for the patient’s vascular status, referral to a vascular surgeon may be necessary.

FIGURE 44.1 Varus heel deformity. (A) Standing examination of the patient from the front reveals a “peek-a-boo” heel due to varus hindfoot. The arrow in the image shows the “peek-a-boo” heel. Note this patient appears to have unilateral cavus of the right foot. (B) Standing exam of the patient from behind shows varus hindfoot alignment. (C) Coleman block testing shows correction of hindfoot varus to neutral.

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44.5

PART | 7 Clincial Interventions

Imaging

44.5.1 X-rays Weight-bearing radiographs of the foot and ankle are important to assess the bony deformity of the cavus foot. Cavus alignment will result in increased calcaneal pitch and increased arch height on the lateral view and stacked metatarsals on the anteroposterior (AP) view. The lateral talometatarsal relationship can be used to assess the severity of first

FIGURE 44.2 Weight bearing X-rays of neutral foot alignment. (A) AP view shows neutral talocalcaneal relationship and no significant stacking of metatarsals. (B) Lateral view shows normal calcaneal pitch and neutral lateral talometatarsal relationship. Compare Figs. 44.2 with 44.3 to see changes related to cavus foot alignment.

FIGURE 44.3 Weight bearing X-rays of a cavus foot. (A) AP view shows change in talocalcaneal relationship and stacked metatarsals. (B) Lateral view shows increased calcaneal pitch (blue lines) and change of the lateral talometatarsal relationship with flexion of the first metatarsal (red lines).

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FIGURE 44.4 Weight bearing CT scan coronal image of the left foot showing varus malalignment through the posterior facet of the subtalar joint.

metatarsal flexion. With normal foot alignment, a line drawn down the axis of the talus should pass down the axis of the first metatarsal on the AP and lateral views. With cavus alignment there will be distinct flexion of the first metatarsal and disruption of this relationship (Figs. 44.2 and 44.3). There will also be changes of the talocalcaneal relationship on the AP view. Ankle images can help to assess if there are any secondary changes in ankle alignment due to the underlying cavus foot alignment. As the axis of the ankle is externally rotated, there will be posterior positioning of the fibula in relation to the tibia.

44.5.2 Weight-bearing CT scan A weight-bearing computed tomography (CT) scan can be used to further assess the cavus foot as bony detail can be assessed more thoroughly. Obtaining a weight-bearing CT scan can help to assess for any underlying deformities of the subtalar joint that may be contributing to hindfoot varus (Fig. 44.4). Advanced imaging with a weight-bearing CT scan can be an integral part of treatment planning and execution.

44.6

Biomechanical changes of pes cavus

Pes cavus results in specific mechanical changes to the foot despite the underlying cause of the deformity. While the driving force behind cavus foot alignment can be related to the forefoot, hindfoot, or both, the consequences of this deformity are similar. Pes cavus results in a stiffer foot than normal alignment. This is due to less subtalar and transverse tarsal joint motion, causing a reduced ability of the foot to absorb impact during the early part of stance phase. This is a result of a more vertically oriented subtalar joint axis and the talar head remaining over the anterior process of the calcaneus. There are substantially higher plantar pressures beneath the metatarsal heads and heel pad due to a decrease in the size of the areas that bear weight [12]. Toe deformities are common with cavus alignment. These can include claw toe deformity, hammer toe deformity, and mallet toe deformity. Power is diminished due to the toes not participating in the toe off phase of gait [12]. Claw toe deformities include extension at the MTP joint with flexion at the interphalangeal (IP) joints. Hammer toe deformities include flexion at the proximal IP joint, while mallet toe deformities include flexion at the distal IP joint.

44.6.1 Hindfoot Hindfoot-driven cavus relates to a fixed varus deformity of the hindfoot. Elevated pitch of the longitudinal axis of the calcaneus is present. Fixed hindfoot varus results in compensatory flexion of the first metatarsal [13]. Hindfoot-driven

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deformities were very common during the era of widespread poliomyelitis however they are seen less frequently now [12]. Elevated calcaneal pitch and hindfoot varus are more commonly encountered as the result of a combined deformity. Coleman block testing will not result in correction of the hindfoot in these fixed varus scenarios.

44.6.2 Forefoot Forefoot-driven cavus is caused by the primary deforming force of a flexed first metatarsal. Flexion of the first metatarsal produces increased arch height. Forefoot-driven hindfoot varus results from the position that the flexible hindfoot assumes when compensating for the flexed first ray [14]. During the foot flat and push off phases of gait, the medial forefoot strikes the ground first which leads to the inability to evert through the subtalar and transverse tarsal joints. The reduction in eversion through the subtalar and transverse tarsal joints creates varus hindfoot alignment and a stiffer foot that does not absorb impact as well as normal alignment. Mosca related the flexed position of the first metatarsal to hyperactivity of the peroneus longus [15]. Manoli et al. coined this “peroneal overdrive” [7]. Initially the flexed position of the first metatarsal is flexible, however with time it becomes stiff and eventually rigid. The hindfoot follows this pattern of progressing from flexible to stiff. With time the resultant deformity involves the entire forefoot developing a valgus position while the hindfoot becomes fixed in varus [7]. Coleman block testing will result in correction of hindfoot varus when the deformity is flexible, however as the deformity becomes more rigid this correction will be lost. Cavus foot alignment is accompanied by various deformities of the toes and MTP joints. Deformities of the toes and MTP joints may be mild involving flexible clawing of the MTP joints with mild flexion of the IP joints but can progress to more severe rigid claw toe deformities [16]. Rigid claw toe deformities can have detrimental effects including displacement of the plantar fat pad distally as the toes pull into extension. The extrinsic toe extensors can also drive the metatarsal heads into a more flexed position, therefore the metatarsal heads are more plantar and without their normal cushioning of the fat pad [17].

44.6.3 Soft tissues Most patients with cavus foot alignment have equinus deformity due to tightness of the gastrocnemius muscle. Gastrocnemius equinus can be evaluated for using the Silverskiold test to isolate the gastrocnemius from the remainder of the triceps surae complex. The Silverskiold test is performed by fully extending the knee and holding the subtalar joint in neutral alignment. A dorsiflexion force is applied symmetrically across the forefoot and the angle of dorsiflexion at the ankle is noted. The knee is then flexed, and the force is repeated. If the patient has no passive ankle dorsiflexion with the knee extended and corrects to approximately 5 degrees or greater with the knee flexed, gastrocnemius equinus exists [18]. Plantarflexion at the ankle that is produced by gastrocnemius equinus increases the mechanical advantage of the peroneus longus compared to its antagonist, the tibialis anterior. Equinus deformity allows for increased flexion of the first metatarsal by the peroneus longus. Ankle dorsiflexion is also impacted by forefoot valgus. Relative forefoot equinus is created with forefoot valgus and the weight bearing portion of the foot is more flexed due to the head of the first metatarsal being plantar to the plantar aspect of the calcaneal tuberosity. Cavus foot alignment and increased arch height cause shortening of the plantar aponeurosis. The plantar aponeurosis typically functions as a windlass mechanism during gait to elevate the arch, flex the metatarsals, and invert the hindfoot [12]. These conditions are already present in the cavus foot and therefore contracture of the plantar aponeurosis can develop. As contracture occurs, the plantar aponeurosis helps to hold the elevated arch height, forefoot internal rotation, and hindfoot inversion [12].

44.7

Conservative management

The goal of treatment including conservative management is to control symptoms by creating a plantigrade foot and to redistribute weight-bearing forces. An important component of conservative treatment is stretching and range of motion exercises to maintain flexibility. If there is gastrocnemius equinus noted on physical examination, an aggressive gastrocnemius stretching protocol should be recommended. Exercises to maintain ankle, subtalar, transverse tarsal, and MTP joint range of motion are also recommended [7]. Inserts and bracing may be used to help control symptoms related to cavus foot alignment. An orthosismay be used to help correct hindfoot varus and offload areas of increased pressure caused by cavus foot alignment [19]. Articulated or solid custom ankle-foot orthoses may be needed to assist in symptom control in more severe cases. Inserts and bracing are used to gain a plantigrade foot and relieve symptoms due to cavus foot alignment.

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44.8

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Surgical management

44.8.1 Overview When nonsurgical treatment fails, a myriad of surgical options exist to correct the cavovarus deformity for achieving a well-balanced, plantigrade foot. Surgical treatment of the cavus foot can be broadly divided into joint sparing procedures for a flexible deformity which use a combination of soft tissue releases, tendon transfers, and osteotomies; and joint sacrificing procedures which achieve correction of severe or arthritic deformities through fusions of joints of the hindfoot and midfoot. Both joint sparing and joint sacrificing procedures can achieve a well-balanced, plantigrade foot when used appropriately. Differentiating between joint sparing and joint sacrificing procedures involves a careful and thorough preoperative evaluation by the surgeon. A comprehensive preoperative evaluation should include a thorough physical exam and detailed review of the patient’s imaging. The soft tissue must be carefully inspected and the incisions marked out prior to surgery to identify potential areas of wound complications. In correction of severe deformities, the surgeon must be cognizant of the potential for wound complications medially. A detailed review of the imaging including weight bearing X-rays of the foot and ankle can identify the extent and location of deformity and degree of arthritis. Advanced imaging including weight bearing CT and magnetic resonance imaging (MRI) scans may also be helpful and provide detailed anatomic images and assess the degree of degenerative changes. Finally, the patient’s age, activity level, and expectations must be taken into account. For most mild-to-moderate deformities that are flexible and have minimal degenerative changes, surgical correction can be achieved with a gastrocnemius recession or percutaneous Achilles tendon lengthening, open plantar fascia release, peroneus longus to brevis transfer, extension osteotomy of the first metatarsal, and lateralizing calcaneal osteotomy. Additional procedures may be necessary and include capsular releases, tendon lengthening, a posterior tibial tendon transfer, tarsal tunnel release and/or corrective osteotomies and fusions in the midfoot. In severe deformities or deformities associated with degenerative changes in the joints of the hindfoot, surgical correction can be achieved with a triple arthrodesis in combination with soft-tissue release, tendon transfers, and osteotomies. Forefoot procedures commonly performed in patients with cavus feet include a modified Jones procedure and Girdlestone-Taylor flexor-toextensor tendon transfers for extensor clawing of the hallux and lesser toes respectively and may be necessary to achieve the patient’s desired outcome. The surgeon must keep in mind that cavus deformities are heterogeneous and a surgical plan must be individualized to the patient’s deformity and presenting complaints. It is also imperative to identify manifestations of the cavus foot such as lateral ankle instability, recurrent Jones fractures, and anteromedial ankle arthritis and treat them if necessary. Often times it is these manifestations of the cavus foot that are the presenting complaint when a patient seeks surgical intervention. When these manifestations are treated independently, without addressing the underlying cavus deformity, there may be an increased rate of recurrence and failure of surgery. Patients should be counseled, and the deformity should be corrected in conjunction with the presenting complaint if necessary.

44.8.2 Gastrocnemius recession/Achilles tendon lengthening Surgical lengthening of the gastrocnemius-soleus complex is an integral part of a cavus foot reconstruction and serves several biomechanical functions. Performing a gastrocnemius recession or Achilles tendon lengthening eliminates the varus moment to the hindfoot resulting from a tight heel cord. This facilitates lateral calcaneal translation of the calcaneal osteotomy. It also reduces forefoot pressure which can help alleviate metatarsalgia from chronically dislocated MTP joints as a result of extensor clawing due to intrinsic muscle weakness. The Silfverskiold test distinguishes whether a gastrocnemius recession or Achilles tendon lengthening is indicated. An isolated contracture of the gastrocnemius can be lengthened at several levels in the leg. The Strayer procedure lengthens the gastrocnemius in the mid-calf just proximal to where the gastrocnemius aponeurosis coalesces with the soleus [20]. A longitudinal incision is made over the posteromedial leg, just distal to the medial muscle belly of the gastrocnemius. The superficial fascia is identified and incised. Blunt dissection allows the gastrocnemius aponeurosis to be dissected away from the superficial fascia and soleus. The sural nerve can be adherent to the gastrocnemius aponeurosis lateral to the midline and should be identified and carefully retracted. Modified right-angle retractors can be used to isolate the gastrocnemius aponeurosis which is released sharply from medial to lateral. If additional dorsiflexion is required, the soleus fascia can be incised deep to the gastrocnemius aponeurosis. For a global Achilles tendon contracture, a percutaneous triple hemi-section of the Achilles tendon can be performed. Three stab incisions are made at 2 cm intervals over the distal Achilles tendon. The medial fibers are released

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through the most proximal and distal stab incisions and the lateral fibers are released in the middle stab incision. A gentle dorsiflexion force is applied to the ankle until the desired correction is achieved. Rarely, an open Z-lengthening may be required.

44.8.3 Plantar fascia release Due to the deforming force of the peroneus longus overpowering the weakened tibialis anterior, the first ray becomes flexed and as a result the plantar fascia becomes contracted. Releasing the plantar fascia helps to correct midfoot cavus and should be done prior to a dorsiflexion osteotomy of the first metatarsal to maximize the amount of correction the osteotomy is able to achieve. A short longitudinal incision is made along the medial heel just distal to the calcaneus. The abductor hallucis fascia is identified and carefully retracted plantarly. The plantar fascia is carefully isolated and sharply released from medial to lateral. The main structure at risk is the medial calcaneal nerve.

44.8.4 Peroneus longus to brevis transfer A peroneus longus to brevis transfer corrects the deforming force of the flexed first metatarsal and provides an eversion force to the hindfoot. The transfer can be performed posterior to the lateral malleolus or over the lateral wall of the calcaneus. To transfer the peroneus longus to brevis posterior to the lateral malleolus, a longitudinal incision is made over the posterior border of the lateral malleolus. The peroneal tendon sheath is identified and incised. The peroneus brevis tendon is identified directly posterior to the fibula. The peroneus longus tendon is identified and released distally. The paratenon of the peroneus brevis is stripped and the peroneus longus is then transferred to the peroneus brevis and secured via a Pulvertaft weave. The tendon transfer is secured with nonabsorbable suture while the hindfoot is everted. If the peroneus brevis is irreparably torn, the peroneus longus can be released distally and secured to the fifth metatarsal base.

44.8.5 Posterior tibial tendon lengthening and transfer One of the hallmarks of a flexible cavovarus deformity seen in CMT is selective weakness of the peroneus brevis which is then overpowered by its antagonist the posterior tibial tendon. If tibialis anterior function is intact with only mild weakness and the posterior tibial tendon is a major deforming force, a Z-lengthening of the posterior tibial tendon can provide controlled correction. This procedure is performed through a medial incision and additional correction can be achieved by releasing the talonavicular joint capsule. When the patient presents with absent or weak dorsiflexion of the ankle due to absent tibialis anterior function, a posterior tibial tendon transfer may be required. A good solution involves an incision transfer through the interosseous membrane although several other techniques have been described. The posterior tibial tendon is released off the navicular, passed through the interosseous membrane in the distal third of the leg, shuttled under the extensor retinaculum, and secured to the lateral cuneiform with a bio-tenodesis screw or bone tunnel. Alternatively, the posterior tibial tendon can be transferred to the peroneus tertius tendon [21].

44.8.6 Dorsiflexion osteotomy or fusion of first ray If the first metatarsal remains rigidly flexed after a peroneus longus to brevis transfer and plantar fascia release, a dorsal closing wedge osteotomy of the first metatarsal or dorsiflexion fusion of the first tarsometatarsal (TMT) joint can provide further correction of a fixed deformity. The plantar fascia release and peroneus longus to brevis transfer should be performed before the osteotomy of the first metatarsal so the maximum amount of correction can be obtained. An incision is made in the interspace between the first and second metatarsals. The extensor hallucis longus (EHL) tendon sheath is incised and the EHL tendon is retracted laterally. The osteotomy is started 1.5 cm distal from the first TMT joint. The size of the dorsal closing wedge varies depending on the amount of correction required, but generally the dorsal wedge should be 4 5 mm. The orientation of the proximal osteotomy should be parallel to the first TMT joint and the more distal cut should be perpendicular to the long axis of the first metatarsal. Both cuts should converge plantarly, leaving the plantar bone intact as a hinge. Gentle force under the first metatarsal closes the dorsal wedge, correcting the flexion deformity. Once the correction is achieved, fixation can be provided with a lag screw, tension band construct, or dorsal plate and screws. Alternatively, the closing wedge can be removed from the first TMT joint and fusion of the first TMT joint can be performed to correct the flexion deformity of the first metatarsal (Figs. 44.4 and 44.5).

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FIGURE 44.5 (A) Lateral view pre-operatively showing flexed first metatarsal. (B) Lateral view postoperatively showing correction through dorsiflexion fusion of the first TMT joint. Courtesy Jeff Henning, MD.

44.8.7 Calcaneal osteotomy Correction of fixed hindfoot varus can be accomplished with a Dwyer closing wedge osteotomy [22], lateralizing calcaneal osteotomy, or a Malerba Z-cut osteotomy [23]. Cadaveric studies have shown that all three osteotomies are able to produce significant lateralization and decrease in peak forces [24]. The Malerba Z-cut osteotomy provides the greatest amount of correction but is the most aggressive and over-correction can occur. Lateralizing calcaneal osteotomy can usually provide adequate correction of both hindfoot varus and increased calcaneal pitch and can be used in the majority of flexible cavus reconstructions cases. While some authors describe performing an isolated closing wedge osteotomy of the first metatarsal for cavus deformities that have correction of hindfoot varus with Coleman block testing (forefoot-driven varus), a lateralizing calcaneal osteotomy is still often indicated for flexible cavus foot reconstructions even if there is correction of hindfoot varus with Coleman block testing. The Malerba Z-cut osteotomy is considered for severe deformities and revisions. It is important to address a tight heel cord prior to performing the calcaneal osteotomy to remove the varus moment arm and facilitate lateral translation of the calcaneal tuberosity. The lateralizing calcaneal osteotomy is usually performed through an oblique incision over the lateral aspect of the calcaneus, although a medial approach has been described [25]. After careful dissection down the lateral wall of the calcaneus, the osteotomy site is confirmed with fluoroscopy and retractors are placed superiorly and inferiorly to the tuberosity. The orientation of the osteotomy is critical and should be perpendicular to the calcaneus or cheated from lateral to posteromedial; if the orientation is made in the opposite direction shifting the tuberosity laterally is much more difficult. A microsagittal saw with a 3 4 cm long blade is used to cut from lateral to medial, taking care not to inadvertently over-penetrate medially, potentially damaging neurovascular structures. After the cut is made the osteotomy site should be distracted with a lamina spreader for 1 2 minutes and a blunt, right angle instrument should be used to release adherent soft tissues medially to facilitate lateral translation. Once the tuberosity is freely mobile it should be translated 6 10 mm laterally for correction of varus and elevated several millimeters for correction of increased calcaneal pitch. The tuberosity is provisionally held with k-wires while the correction is verified with fluoroscopy using an axial heel view. The osteotomy is secured with two partially threaded 4.5 or 6.5 mm screws. Recent evidence suggests similar rates of healing but lower rates of hardware removal when 4.5 mm screws are used [26]. Alternatively, 3.5 mm cortical screws placed in lag fashion for compression can be used to secure the osteotomy site. After the osteotomy is secured, the lateral edge of the posterior tuberosity is beveled to prevent a painful prominence (Fig. 44.6). A potential complication of the lateralizing calcaneal osteotomy is an acute tarsal tunnel syndrome. As the tuberosity is shifted laterally there is a decrease in volume of the tarsal tunnel which may cause compression of the posterior tibial nerve [27]. Up to 34% of patients may have tibial nerve symptoms following a lateral calcaneal osteotomy, but the majority seem to resolve with time and prophylactic tarsal tunnel release does not alter the incidence [28]. There is little need to routinely perform a prophylactic tarsal tunnel release but the possibility of transient or permanent tibial nerve symptoms postoperatively should be discussed with the patient. If an aggressive lateral shift is planned a prophylactic tarsal tunnel release may be considered.

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FIGURE 44.6 (A, B) AP and lateral views pre-operatively showing cavus foot alignment. (C, D) AP and lateral views postoperatively showing good correction with a dorsiflexion fusion through the first TMT joint and a lateralizing calcaneal osteotomy. Images courtesy of Chad Carlson, MD.

44.8.8 Modified jones procedure Intrinsic muscle weakness and extensor recruitment causes extensor clawing of the hallux which can cause discomfort from dorsal impingement in the shoe and pain under the plantar aspect of the first MTP joint. Depending on the degree of discomfort caused by extensor clawing and the number of additional procedures being performed, a modified Jones procedure can be performed for correction at the same time as the cavus reconstruction. If too many procedures are being performed to safely add on forefoot procedures, the hallux and lesser toes can be staged and addressed at a later surgical procedure. The modified Jones procedure includes a fusion of the hallux IP joint and transfer of the EHL tendon to the first metatarsal neck. The hallux IP joint is approached dorsally and flat cuts are made with a small microsagittal saw through the distal portion of the proximal phalanx and the base of the distal phalanx. The joint is provisionally held with a k-wire and a partially threaded 4.5-mm screw or a 4.0 mm cortical lag screw is placed through a stab incision in the tip of the toe across the fusion site. A temporary k-wire can be placed across the fusion site for rotational stability and removed in the office after 6-weeks. After the hallux IP joint is secured, the EHL tendon is released off the distal phalanx, passed through a drill hole oriented from medial to lateral in the neck of the first metatarsal and secured to itself with a nonabsorbable suture. This allows the EHL to continue to assist in dorsiflexion of the ankle and resist against flexion of the first metatarsal.

44.8.9 Lesser toe deformities Flexible or rigid claw toes may develop as the intrinsic muscles of the foot are overpowered by the extrinsic muscles of the leg. Depending on the number of procedures being performed and the comfort level of the treating surgeon, symptomatic claw toe deformities can be corrected at the time of the cavus foot reconstruction or as a separate surgical procedure. Surgical correction of claw toes depends on whether the toe is flexible or rigid. In flexible claw-toes a Girdlestone-Taylor flexor-to-extensor tendon transfer is performed. In rigid deformities, correction is achieved through a proximal IP joint resection arthroplasty and Girdlestone-Taylor flexor-to-extensor tendon transfers. If the MTP joints are subluxated or dislocated a combination of extensor tendon lengthening, dorsal MTP capsulotomy, and metatarsal shortening osteotomy may be required for correction and reduction of the MTP joint. If a metatarsal shortening osteotomy is indicated, a midshaft shortening osteotomy is preferable. The midshaft osteotomy allows controlled shortening. The osteotomy is secured with a 4-hole quarter tubular plate applied dorsally over the metatarsal. Placing an apex dorsal bend in the plate and placing the screws eccentrically generates compression across the osteotomy site. Bone graft may be required to fill in any gaps.

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44.8.10 Triple arthrodesis and midfoot fusion For severe deformities that are not passively correctable or for deformities where arthritic changes are present, full correction may require fusions of the hindfoot or midfoot depending on the location of the deformity and arthritic changes. Full correction of the cavus deformity may require additional tendon transfers and osteotomies in addition to the correction achieved through a triple arthrodesis to achieve a plantigrade foot. Similar to correction of the flexible cavus deformity, soft-tissue releases should be performed as needed to allow maximum correction. Isolated fusion of the subtalar joint or combined fusion of the talonavicular, subtalar, and calcaneocuboid joints can provide powerful correction of the cavus deformity. Dual incisions both medially and laterally are required for preparation of the joint surfaces. A lateral incision from the tip of the lateral malleolus extending towards the base of the fourth metatarsal allows exposure of the subtalar joint, calcaneocuboid joint, and lateral aspect of the talonavicular joint. The majority of the talonavicular joint is approached through a medial incision from the medial malleolus extending down the medial column of the foot. Meticulous preparation of the joint surfaces is necessary to remove remaining articular cartilage and prepare the subchondral bone for fusion. A bur can be used to remove bone to allow maximum correction and bone graft may be needed to fill voids resulting from the deformity. The hindfoot joints can be secured with solid, fully and partially threaded screws, as well as cannulated screws, plate and screw constructs, and staples. The surgeon should be prepared to perform a lateralizing calcaneal osteotomy or closing wedge dorsiflexion osteotomy after the triple arthrodesis if further correction is needed (Fig. 44.7).

FIGURE 44.7 (A) Lateral view pre-operatively showing cavus foot alignment with increased calcaneal pitch and flexed first metatarsal. (B, C) Sagittal and coronal views of a CT scan showing degenerative changes of the subtalar joint and varus malalignment of the posterior facet of the subtalar joint. (D) Lateral view postoperatively showing correction with a subtalar fusion, calcaneocuboid fusion, and dorsiflexion fusion through the first TMT joint.

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If the apex of the deformity is in the midfoot and involves multiple metatarsals, a more extensive midfoot osteotomy may be required. Several different osteotomies have been described including the Jahss truncated wedge osteotomy [29], Cole midtarsal osteotomy [30], and Japas midtarsal osteotomy [31]. Additional tendon transfers, osteotomies, and soft-tissue releases may be required for complete deformity correction.

44.8.11 Treatment of associated ankle pathology With chronic cases of cavus foot alignment, varus ankle instability, and ankle arthritis with varus ankle alignment, additional surgical procedures beyond the cavus reconstruction may be warranted to treat the associated ankle pathology. Procedures could include supramalleolar tibial osteotomies to realign weight-bearing through the lateral portion of the ankle. Ankle fusion or ankle replacement may be warranted with severe degenerative changes of the tibiotalar joint (Fig. 44.8).

44.8.12 Postoperative care Despite the wide variety of surgical procedures used to correct the cavus foot, postoperative treatment is fairly consistent. The patient is placed in a postoperative bulky-Jones splint for the first 2 weeks to allow surgical incisions to heal and to protect the surgical reconstruction. The splint and sutures are removed at the first postoperative visit and the patient is transitioned to a nonweight bearing short leg cast for the next 4 6 weeks. At 6 8 weeks partial weight bearing is initiated in a tall controlled ankle motion boot and the patient is progressed to full weight bearing over the next month. The patient is then weaned out of the boot and referred to physical therapy. The patient is encouraged to wear a comfortable shoe and inserts and an ankle brace for high risk activities. The patient is counseled that full recovery can take up to one year.

FIGURE 44.8 (A) Lateral view pre-operatively showing cavus foot alignment and degenerative changes of the tibiotalar joint. (B) AP view of the ankle pre-operatively showing severe varus alignment of the talus. (C, D) Lateral and AP view postoperatively showing ankle fusion with correction of the underlying cavus deformity with a lateralizing calcaneal osteotomy and dorsiflexion fusion of the first TMT joint. Courtesy Ben Stevens, MD.

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44.8.13 Complications In addition to the surgical risks inherent in any lower-extremity surgery, there are several complications particularly relevant to surgical correction of the cavus foot. Under correction of the deformity can occur and is best prevented by careful preoperative planning to fully understand the deformity and what combination of surgical procedures are required for a full correction. Frequently reassessment of the correction during surgery is also necessary to confirm that complete correction has been achieved. A failure to address the underlying muscle imbalances leading to a flexible deformity can also contribute to recurrence. As mentioned previously, an aggressive lateralizing calcaneal osteotomy can reduce the volume of the tarsal tunnel and lead to postoperative tibial nerve symptoms. In most patients, this postoperative tarsal tunnel syndrome is transient and not affected by a prophylactic tarsal tunnel release [27]. As the deformity is corrected, the contracted medial soft tissues can put tension on incisions, increasing the risk for delayed wound healing and wound necrosis. Careful preoperative planning of the incisions, meticulous soft tissue handling, and a tension free closure can help to minimize medial soft tissue complications.

44.9

Areas of future research

Treatment of varus ankle arthritis in the setting of a cavus foot is a challenging clinical problem with no clear consensus on the best treatment. The cavovarus deformity directly contributes to the development of arthritis due to increased medial joint pressures and may also lead to recurrent ankle sprains due to varus ankle instability [32,33]. The cavus foot and lateral ankle instability can lead to a spectrum of degenerative changes from isolated anteromedial ankle impingement to end-stage varus ankle arthritis. In the younger patient with mild to moderate arthritis, correction of the deformity and an ankle cheilectomy can realign the foot and prevent progression of arthritis. In more advanced ankle arthritis with increasing degrees of varus deformity surgical decision making can be more difficult. Historically, an ankle fusion or tibiotalocalcaneal fusion was required to address both the deformity and ankle arthritis. Varus arthritis of over 10 degrees was traditionally considered to be a contraindication for ankle replacement due to early failure of the implants. As the total ankle implants and indications have evolved, increasing varus deformity in the ankle may be tolerated with newer generation implants and concomitant correction of the deformity. Further research is required on the expanding role of total replacements in the treatment of more severe, end-stage varus ankle arthritis.

References [1] Sangeorzan BJ, Wagner UA, Harrington RM, Tencer AF. Contact characteristics of the subtlar joint: the effect of talar neck malalignment. J Orthop Res 1992;10:544 51. [2] Brewerton DA, Sandifer PH, Sweetnam DR. Idiopathic pes cavus: an investigation into its aetiology. BMJ 1963;2:659 61. [3] Holmes JR, Hansen ST. Foot and ankle manifestations of Charcot-Marie-tooth disease. Foot Ankle 1993;14:476 86. [4] Tynan MC, Klenerman L, Helliwell TR, Edwards RH, Hayward M. Investigation of muscle imbalance in the leg in symptomatic forefoot pes cavus: a multidisciplinary study. Foot Ankle 1992;13:489 501. [5] Bertorini T, Narayanaswami P, Rashed H. Charcot-Marie-Tooth disease (hereditary motor sensory neuropathies) and hereditary sensory and autonomic neuropathies. Neurologist 2004;10(6):327 37. [6] Morcuende JA, Dolan LA, Dietz FR, Ponseti IV. Radical reduction in the rate of extensive corrective surgery for clubfoot using the Ponseti method. Pediatrics 2004;113(2):376 80. [7] Manoli A, Graham B, Ped C. The subtle cavus foot, “The Underpronator,” a review. Foot Ankle Int 2005;26(3):256 63. [8] Alexander IJ, Johnson KA. Assessment and management of pes cavus in Charcot-Marie-tooth disease. Clin Orthop 1989;246:273 81. [9] Saxena A, Krisdakumtorn T, Erickson S. Proximal fourth metatarsal fractures in athletes: similarity to proximal fifth metatarsal injury. Foot Ankle Int 2001;22:603 8. [10] Younger AS, Hansen ST. Adult cavovarus foot. J Am Acad Orthop Surg 2005;13:302 15. [11] Brandes CB, Smith RW. Characterization of patients with primary peroneus longus tendinopathy: a review of twenty-two cases. Foot Ankle Int 2000;21:462 8. [12] Guyton GP, Mann RA. Pes Cavus. In: Coughlin MJ, et al., editors. Surgery of the foot and ankle, vol. 1. Mosby Elsevier; 2007. p. 1125 48. [13] Solis G, Hennessy M, Saxby TS. Pes cavus: a review. Foot Ankle Surg 2000;6(3):145 53. [14] Apostle KL, Sangeorzan BJ. Anatomy of the varus foot and ankle. Foot Ankle Clin 2012;17(1):1 11. [15] Mosca VS. The cavus foot. J Pediatr Orthop 2001;21:423 4. [16] McCluskey WP, Lovell WW, Cummings RJ. The cavovarus foot deformity. Etiology and management. Clin Orthop 1989;247:27 37. [17] Heyman CH. The operative treatment of clawfoot. J Bone Jt Surg 1932;14:335 8. [18] DiGiovanni CW, Kuo R, Tejwani N, et al. Isolated gastrocnemius tightness. J Bone Jt Surg 2002;84A:962 70. [19] Bordelon RL. Orthotics, shoes, and braces. Orthop Clin North Am 1989;20:751 7.

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[20] Pinney SJ, Sangeorzan BJ, Hansen ST. Surgical anatomy of the gastrocnemius recession (Strayer procedure). Foot Ankle Int 2004;25 (4):247 50. [21] Hsu J, Hoffer M. Posterior tibial tendon transfer anteriorly through the interosseous membrane: a modification of the technique. Clin Orthop Relat Res 1978;131:202 4. [22] Dwyer FC. Osteotomy of the calcaneum for pes cavus. J Bone Jt Surg (Br) 1959;41b(1):80 6. [23] Malerba F, De Marchi F. Calcaneal osteotomies. Foot Ankle Clin 2005;10(3):523 40. [24] Krause FG, et al. Ankle joint pressure changes in a pes cavovarus model after lateralizing calcaneal osteotomies. Foot Ankle Int 2010;31 (9):741 6. [25] Jaffe D, Vier D, Kane J, Kozanek M, Royer C. Rate of neurologic injury following lateralizing calcaneal osteotomy performed through a medial approach. Foot Ankle Int 2017;38(12):1367 73. [26] Lucas ED, et al. Screw size and insertion technique compared with removal rates for calcaneal displacement osteotomy. Foot Ankle Int 2015;36 (4):395 9. [27] Bruce BG, Barieau JT, Evangelista PE, Arcuri D, Sandusky M, DiGiovanni CW. The effect of medial and lateral calcaneal osteotomies on the tarsal tunnel. Foot Ankle Int 2014;35(4):383 8. [28] Van Valkenburg S, Hsu RY, Palmer DS, Blankenhorn B, Den Hartog BD, DiGiovanni CW. Neurologic deficit associated with lateralizing calcaneal osteotomy for cavovarus foot correction. Foot Ankle Int 2016;37(10):1106 12. [29] Jahss MH. Tarsometatarsal truncated-wedge arthrodesis for pes cavus and equinovarus deformity of the fore part of the foot. J Bone Jt Surg (Am) 1980;62:713 22. [30] Cole WH. The treatment of claw foot. J Bone Jt Surg 1940;22:895 908. [31] Japas LM. Surgical treatment of pes cavus by tarsal V-osteotomy. J Bone Jt Surg (Am) 1968;50:927 44. [32] Krause F, Windolf M, Schweiger K, Weber M. Ankle joint pressure in pes cavovarus. J Bone Jt Surg Br 2007;89:1660 5. [33] Valderranbano V, Hinterman B, Horisberger M, Fung TS. Ligamentous post-traumatic ankle osteoarthritis. Am J Sports Med 2006;34:612 20.

Chapter 45

Biomechanics of Hindfoot Fusions Dante Marconi1 and Andrew K. Sands2 1

Kingsbrook Jewish Medical Center, Shore Physicians Group, Somers Point, Brooklyn, NY, United States, 2New York Downtown Orthopaedic

Associates, New York, NY, United States

Abstract The hindfoot consists of the talus, calcaneus, navicular, cuboid, and all of the associated articulations. Movement is complex among these joints and is crucial to maintain normal gait. The hindfoot is unlocked during heel strike and midstance phase of gait to accommodate uneven surfaces, and it is locked during toe-off to provide a stable lever for push-off. Deformity of these joints is common, and diagnosis can be made through physical exam and radiographs; however, sometimes advanced imaging is necessary. Foot deformities, such as flatfoot and cavus foot, contribute to the majority of cases of hindfoot pathology. Trauma to this area can also create malalignment and/or arthritis that requires correction. Sequelae of hindfoot pathology can lead to gait instability, ankle sprains, tendinopathies, and can also cause symptoms in other joints such as the knee, hip, and back. Many treatments to address dysfunction have evolved throughout the years. The main goal is to maintain maximal motion that also allows stable gait with minimal pain and symptoms. For mild cases, shoe wear modifications, inserts, and orthotics, as well as activity modification may be appropriate. For the more severe cases of hindfoot pathology, arthrodesis may be the best option. The type of arthrodesis depends on which joints are involved. Some examples include subtalar and talonavicular arthrodesis, triple arthrodesis, and pantalar arthrodesis if the ankle joint is involved.

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Introduction

In the higher primate, the anthropoid, the foot was used for grasping and mobility. The function of the foot has since evolved to include bipedal gait and stability [1]. Certain motion segments are required for functions such as accommodation to uneven surfaces and shock absorption. These essential joints are defined as those that should not be sacrificed to maintain normal function. These include the tibiotalar, the talocalcaneal (which includes three facets), the talonavicular, the lesser metatarsophalangeal, and the fourth and fifth tarsometatarsal joints. Motion at the talonavicular joint is central to maintaining motion at other joints [2]. Other joints are nonessential stability joints, such as the calcaneocuboid joint in the hindfoot. Other nonessential joints include the intertarsal joints and the first, second, and third tarsometatarsal joints [1]. However, if the essential joints become misaligned, arthritic and/or dysfunctional, this will affect the ability to stand upright, and also will affect the gait cycle. Nonoperative treatment options exist for mild to moderate pathology, but in advanced hindfoot pathology, select realignment arthrodesis of the hindfoot is required. In 1923, Ryerson discussed deformities created during World War I which were only accommodated and not corrected [3]. Shoemakers would design custom shoes, but the deformity persisted. Later, tenotomies and forcible corrections were used to address congenital club foot, but this did not create a permanent cure. After this, tendon transfers were utilized with some success, especially in association with osteotomies. Whitman performed an astragalectomy/ talectomy, with “backward displacement” of the foot to address calcaneus and calcaneal-valgus deformities [3]. The first hindfoot arthrodesis procedures were originally performed in children with polio, Charcot-Marie-Tooth disease, cerebral palsy, and recalcitrant clubfeet [4]. This was after it became evident that patients with a prosthetic leg were better suited for ambulation than patients with infantile paralysis [3]. The artificial leg limited motion in flexion and extension and did not have the pain and/or instability associated with lateral motion. This led to the theory of reconstructing the native bones of the foot to prevent any lateral deformity or instability [3]. In 1913, Gwilym G. Davis Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00032-9 © 2023 Elsevier Inc. All rights reserved.

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described the concept of simultaneous fusion of the talocalcaneal and the talonavicular joints for correcting varusvalgus and abduction-adduction deformities of the foot [5]. It is thought that Dr. Ryerson added the third fusion, the calcaneocuboid joint, creating the triple arthrodesis to further stabilize the hindfoot [3]. Failure to fuse the calcaneocuboid and tarsometatarsal joints were the main causes of recurrent deformity, and over time, including the calcaneocuboid in the fusion gained acceptance, but including the tarsometatarsal joints were abandoned. In 1949, Glissan developed his principles of arthrodesis. These were as follows: adequate joint debridement and preparation, accurate coaptation of surfaces, optimal position, and maintenance of position until arthrodesis is sound [6].

45.2

Complex hindfoot biomechanics

During walking, the normal foot and ankle work in tandem to produce a net mechanical work close to zero [7]. The mechanical work is produced through elastic and viscoelastic mechanisms as well as through muscle contractions [7]. Elastic mechanisms include the Achilles tendon and plantar fascia. An example of a viscoelastic mechanism is the heel pad. The plantar aponeurosis transmits large forces between the hindfoot and the forefoot during gait [8]. Tension of the plantar aponeurosis is low at heel strike, increases during midstance, and peaks at about 80% of the stance phase [8]. Surgical procedures that release the plantar aponeurosis (to decrease arch height) may affect efficient propulsion [8]. The subtalar joint (STJ) consists of an anterior and a posterior part (Fig. 45.1). Anteriorly, the talar head is located on the anterior and middle facets of the calcaneus. Both of these facets are concave. The talar head is also supported by the spring ligament and navicular, and insufficiency of the spring ligament contributes to flatfoot deformity [9]. Posteriorly, the posterior facet of the calcaneus is convex. This facet is larger than the anterior and the middle facets, and it is separated by the interosseous talocalcaneal ligament. Malfunction of the interosseous talocalcaneal ligament can lead to abnormal anterolateral rotation of the talus during gait and cause instability. This problem is accentuated with disruption of the anterior talofibular ligament. The importance of the calcaneofibular ligament in stability is debatable. The inferior extensor retinaculum is important for stability to provide proper gliding of the extensor tendons. The function of the STJ is to cushion heel-strike and to provide stabilization of the midfoot for toe-off during the gait cycle. It protects the proximal joints from impact overload and secondary arthrosis. STJ motion is insignificant for gait and posture, because its lever arms are short [10]. Internal rotation of the tibia causes eversion of the hindfoot and “softening” of the midfoot at heel-strike; this allows for adaptation of the hindfoot and midfoot to uneven surfaces by inversion or eversion. External rotation of the tibia causes inversion of the hindfoot and increased rigidity through the midfoot for leverage at toe-off. These are passive but important functions to protect the ankle from abnormal stresses. A valgus hindfoot increases talonavicular and calcaneocuboid motion. Conversely, a varus hindfoot restricts this motion (Fig. 45.2). Lateral peritalar subluxation and posterior tendon dysfunction leads to sinus tarsi impingement and pain which subsequently leads to lateral tilt and ankle arthrosis. In flatfoot deformity, the hindfoot is usually mobile and in a cavus foot, the hindfoot is usually stiff.

FIGURE 45.1 (A) Simplified graphic of the subtalar joint that glides via these convex and concave facets (superolateral oblique view of left foot). (B) Representative tals The blue portion depicts the talus and the red portion portrays the calcaneus. (B) Anatomical graphic of the subtalar joint demonstrating the talus (blue) and the calcaneus (red).

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Talonavicular Axis Talus

Calcaneocuboid Axis

Lat.

Med.

Calcaneus

Neutral

Inverted Calcaneus

FIGURE 45.2 Coronal plane image depicting the talonavicular, calcaneocuboid, and talocalcaneal joints. When the calcaneus is inverted, the talonavicular and calcaneocuboid axes are convergent, and therefore, locked. Blackwood CB, Yuen TJ, Sangeorzan BJ, Ledoux WR. The Midtarsal Joint Locking Mechanism. Foot Ankle Int. 2005 Dec;26(12):1074 80. The original source of photo was from Elftman H. The transverse tarsal joint and its control. Clin. Orthop. 16:41 45, 1960, with permission for Blackwood publication.

A healthy STJ can compensate for supramalleolar deformities, keeping the hindfoot aligned. Inversion range of motion is greater than eversion, which helps to explain why compensation is greater in a valgus ankle. The STJ is normally positioned in about five degrees of valgus during weightbearing, which is likely the reason why less compensation is required in a varus ankle (Fig. 45.3) [9]. Alteration in the normal architecture of the calcaneus will change its function as a lever arm, as a vertical support, and as a horizontal support of the lateral column. As a lever arm, the tuber acts to translate the triceps surae contracture through a fulcrum in the mid-talus to the forefoot. As a vertical support, the calcaneus must be normal height to maintain length and alignment directly under the tibia to avoid tilt stresses in the ankle. The lateral column support is maintained by the horizontal length of the calcaneus, which controls adduction and abduction of the forefoot. A short lateral column causes lateral peritalar subluxation and reduces the push-off efficiency while also overloading the posterior tibial tendon (PTT) (Fig. 45.4). Adequate force generation from the posterior tibial muscle is necessary for normal foot function. The talonavicular joint resists compressive axial loading when the joint is in the horizontal position [10]. Motion at the calcaneocuboid joint is restricted by bone shape, more specifically, the anterior process of the calcaneus and the plantar medial process of the cuboid [11]. This shape, and also the plantar ligaments, protects the foot from intrinsic midtarsal extension. It functions as an “expansion joint” in the subtalar complex that is most notable at extremes of motion. When fused with normal or increased lateral length, it does not affect the functional motion of the STJ significantly [12]. The medial longitudinal arch (MLA) is formed by the calcaneus, talus, navicular, cuneiforms, and the first and second metatarsals. The elasticity of this arch is important to provide cushioning and support of the body during motion. The lateral longitudinal arch is formed by the calcaneus, cuboid, and the fourth and fifth metatarsals [13]. The PTT is the most important dynamic stabilizer to the MLA and hindfoot. It inverts the foot and plantarflexes the ankle. During normal gait, the PTT pulls the hindfoot into varus and locks the transverse tarsal joints (i.e., the calcaneocuboid and talonavicular joints) to provide a rigid foot for toe-off. With dysfunction of the PTT, the unopposed peroneus brevis leads to inversion of the forefoot. The calcaneus shifts into valgus (eversion) due to shortening of the lateral hindfoot, and the talus moves into plantarflexion. As collapse progresses, the stress is applied to the secondary structures, such as the spring ligament, the deltoid ligament complex, the plantar fascia, and the Achilles tendon [13]. The spring ligament, also known as the plantar calcaneonavicular ligament, is supported medially by anterior fibers of the deltoid ligament and superficially covered by the PTT. It forms a sling that supports the medial and plantar portions of the talar head and holds the talus in the correct position. If the spring ligament is torn, then the talar head

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(A)

(B)

(C)

FHL tendon

posterior syndesmosis

posterior tibial nerve posterior tibial artery

lig. talo-fibulare posterius lig. calcaneo-fibulare

lig. deltoideum tibialis posterior tendon

peroneal tendons

FDL tendon

achilles tendon FHL = flexor hallucis longus FDL = flexor digitorum longus lig. = ligamentum

FIGURE 45.3 Illustration demonstrating subtalar compensation. (A) Normal ankle alignment. (B) Tibiotalar varus without compensation. (C) Subtalar joint realigned to match tibial axis. This article is distributed under the terms of the Creative Commons Attribution-Non Commercial 4.0 International (CC BY-NC 4.0) license (https://creativecommons.org/licenses/by-nc/4.0/) which permits non-commercial use, reproduction and distribution of the work without further permission provided the original work is attributed.

FIGURE 45.4 Lateral and anterior/posterior radiographs of the foot showing peritalar subluxation. The lateral view shows the lateral talometatarsal (Meary’s) angle which helps to evaluate flat or cavus foot types. Anterior/posterior view shows the transverse talometatarsal angle and talonavicular uncoverage.

rotates inferiorly (i.e., plantarflexes) and medially (i.e., internally rotates), and the calcaneus progressively tilts into valgus [13]. The deltoid medial collateral ligament has superficial and deep layers. It provides restraint against valgus tilt and external rotation of the talus [13]. The tarsal sinus is cone-shaped, and it maintains alignment between the talus and the calcaneus by limiting talar tilting and calcaneal inversion. During eversion, the sinus tarsi closes due to the anterior movement of the lateral

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process of the talus. During inversion, the sinus tarsi opens. This is important to consider during hindfoot surgery. When the patient is supine, and the surgeon cannot visualize the hindfoot alignment, a closed sinus tarsi means the STJ is everted, whereas an opened sinus tarsi means that the STJ is inverted. In certain procedures (tibiotalocalcaneal fusion), it is advisable to set the STJ in slight valgus, as opposed to temporarily moving the STJ in varus for other procedures such as STJ arthroereisis for flatfoot [13]. The final alignment however should never have the hindfoot in varus. The best way in the operating room to visually gage hindfoot alignment is to lift the leg up and look at the heel posteriorly.

45.3

Conditions that may require hindfoot fusion

45.3.1 Flatfoot The mechanics of flatfoot are not well understood. The characteristic pattern of flatfoot deformity presents as calcaneus eversion and talonavicular dorsolateral subluxation with medial displacement of the talar head. Distortion is seen through the talonavicular and calcaneocuboid joints. An effective lengthening of the medial column develops due to the collapse of the medial arch. This causes the lateral column to shorten as it accommodates for this collapse. The incongruous STJ can lead to fibular impingement. Hindfoot arthrosis, sinus tarsi impingement, and other secondary effects can also be seen at the ankle. In flexible flatfoot deformity, loss of the MLA is normal in infants/children due to fat underneath the midfoot. The normal arch should form by eight years of age. Most cases are asymptomatic if persisting into adulthood. The most common cause of fixed flatfoot is tarsal coalition. This occurs in about 1% of the population with about 50% 80% of cases being bilateral [13]. In acquired flatfoot, the MLA is initially normal, but fails due to injury to the supporting structures. The most common cause is increased tendinopathy of the PTT. This can lead to a cascade of structural and functional changes described here due to an imbalance of biomechanical forces that maintain the MLA. Other causes include trauma, arthritic processes, and neuropathic arthropathy [13]. The presence of gastrocnemius equinus also leads to later pathology. Many signs and symptoms can occur in flat foot syndromes. These are due to the medial-sided tension and lateralsided compression. Medial soft tissues structures may show inflammatory changes [10]. The talar head may protrude medially with the navicular abducted [10]. Medial keratotic lesions or in extreme cases even plantar medial ulcers can become apparent [10]. Flat foot can lead to a series of events starting with internal rotation of the tibia due to the increase in the calcaneal eversion angle. More proximally, these alterations in mechanics can lead to patellofemoral malalignment, hip malrotation, pelvic malalignment, and spinal malalignment [14]. Lateral bony compression leads to sinus tarsi pain. The lateral process of the talus slides down the slope of the posterior facet and impacts the floor of the sinus tarsi which blocks further motion leading to lateral hindfoot sinus tarsi pain [10].

45.3.2 Cavus foot syndromes Cavovarus deformity is typically due to idiopathic or neuropathic causes. Malalignment can lead to recurrent sprains, chronic lateral ankle instability, and anteromedial ankle arthritis. Varus hindfoot can occur in isolation or as part of a cavovarus deformity [15]. Cavovarus leads to medial loading and anteromedial cartilage wear with subsequent talar tilt exacerbating the varus deformity. The Achilles tendon insertion is medialized which causes the tendon to pull the hindfoot into inversion. This puts more stress on the lateral structures. The peroneus brevis is usually weak or degenerative from recurrent sprains or a neuropathic condition, and it may no longer be able to help stabilize the lateral ankle causing repeated lateral ankle sprains. Overactivity of the peroneus longus creates a plantarflexed first metatarsal. The dorsiflexion force of the tibialis anterior is decreased which also worsens the deformity. This locks the hindfoot during heel strike which causes increased stresses to the hindfoot, and other parts of the foot. These increased stresses can lead to metatarsalgia, plantar fasciitis, overload of the lateral border of the foot, and lateral ankle sprains. Neuropathic conditions lead to overpull of the peroneus longus, tibialis posterior, and triceps surae muscles due to weakness of the peroneus brevis and tibialis anterior. Weakness of the peroneal muscles is compensated for by recruitment of the long extensors [10]. Stress on the peroneus brevis during heel strike may lead to degeneration or rupture of the tendon which occurs if the calcaneal tuberosity is medial to the long axis of the tibia [10].

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45.3.3 Rheumatoid arthritis In patients with rheumatoid arthritis, the hindfoot is involved in about two thirds of cases [16]. Pes planovalgus is the most common hindfoot deformity. A study by Keenan et al. demonstrated that valgus hindfoot deformity is seen in patients with longer disease duration [16]. These authors found this difference in patients who had been diagnosed with rheumatoid arthritis for 25 years vs patients who were diagnosed for 14 years. Also, these patients had more overall pain, and more forefoot pain compared to patients with normal alignment [16]. Cavovarus deformity can be seen in patients with juvenile-onset rheumatoid arthritis. On clinical exam, it is important to evaluate the weightbearing hindfoot alignment and the longitudinal arch integrity. Cartilage damage, capsular and ligamentous laxity, and/or tendon ruptures can lead to subtalar and lateral peritalar subluxation. Increased stress to the deltoid ligament can lead to talar tilt which exacerbates the hindfoot valgus deformity. Subfibular impingement can occur and can lead to ankle valgus tilt and fibular stress fracture.

45.3.4 Osteoarthritis While most literature focuses on end-stage disease, limited evidence exists on the initial pathology of hindfoot arthritis. Both the metabolic activity of cartilage and the surrounding mechanical environment are thought to play a large role in the disease. Adequate blood supply and stability should be maintained [17]. The articular surface experiences increased loads when the hyaline cartilage starts to thin. As a person ages, the body has a decreased ability to heal microtrauma to the cartilage. A positive feedback cycle leads to continuous degeneration. In the setting of post-traumatic arthritis, the altered alignment leads to changes in the normal mechanical stresses which further accentuates the disease process.

45.3.5 Calcaneal fractures Fractures of the calcaneus and sequelae can lead to subsequent hindfoot arthritis and malalignment. Xu et al. in their cadaveric study showed that an increase in calcaneal width ( . 6 mm) limits subtalar motion and increases contact area [18]. However, this increases the high-pressure contact zone, causes tendinous impingement, and impinges on the lateral malleolus. These changes can lead to abnormal and uncomfortable shoe wear. Reductions in calcaneal length ( . 3 mm) limited external rotation and plantarflexion of the STJ. Reductions in height ( . 3 mm) mainly resulted in subtalar rotatory instability and the overall contact area being decreased, but contact pressures remained almost constant [18].

45.3.6 Talar fractures and dislocations The talus contains three joints which cover .60% of the talus with articular surface. Its vulnerable blood supply portends a high risk of avascular necrosis with displaced fractures. Subtalar fracture/dislocation more commonly occurs medially in about 70% of injuries. However, lateral displacement of the foot causes the most harm to the posterior tibial neurovascular structures. It is associated with osteochondral lesions in the sinus tarsi in injuries due to twisting and "hyperpronation" (i.e., increased eversiosn) [1]. The STJ helps accommodate to uneven ground which prevents the ankle joint from tilting. Ankle tilt is not well tolerated [1]. Injuries to the talar head due to a shear force can lead to dislocation of the talonavicular and calcaneocuboid joints [1]. There may also be fractures of the dorsal navicular as well as the distal talus.

45.3.7 Tarsal coalitions Tarsal coalitions may also contribute to misalignment and deformity. The most common coalition is between the anterior calcaneus and the navicular. The next most frequent coalition occurs between the talus and sustentaculum tali. These two together account for 90% of all coalitions [13]. These are associated with rigid flatfoot due to valgus deformity at the STJ [10]. Tarsal coalitions can usually be diagnosed on plain radiograph alone, although an oblique orientation can be helpful. A computed tomography (CT) scan or magnetic resonance imaging (MRI) scan can be obtained in difficult cases and to further define the coalition in three dimensions for preoperative resection planning. A talar beak prominence at the dorsolateral aspect of the talar head leads to impaired subtalar motion. If this finding is present on the radiograph, a diagnosis of coalition should be considered.

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45.3.8 Accessory navicular An accessory navicular can be seen in about 4% 14% of the population [19]. Three types exist, but only type II becomes symptomatic. The PTT attaches onto the accessory navicular which leads to the PTT being displaced upward and inward altering the normal biomechanics of the PTT which no longer functions as an invertor of the foot [13].

45.4

Presurgical assessment

45.4.1 Clinical exam The patient history and physical exam is very important in evaluating hindfoot pathology. Any previous injuries or surgeries should be elicited. Does the patient have any disability in daily activities and sports? Is the patient impaired by the pain? What previous nonoperative measures have been tried? On physical exam, the hindfoot alignment is observed while standing. The patient’s ankle and subtalar range of motion is evaluated while they are seated at the end of the exam table dangling their legs. Ankle stability should be assessed with the ankle in slight plantarflexion and inversion. An anterior drawer test stresses the anterior talofibular ligament. The presence of anterior instability may be assessed. It should be compared to the contralateral side as some anterior laxity may be normal. The presence of excessive drawer with discomfort or a frank “clunk” is pathologic. Subtalar motion can be restricted, for example in osteoarthritis, or increased, as in patients with peritalar instability. Diagnosis of flatfoot is typically done clinically. These patients usually complain about pain in the medial aspect of the ankle. In later stages, pain in the lateral ankle can become apparent due to impingement from the valgus position of the hindfoot. When evaluating the patient from behind, the “too many toes” sign can be seen due to midfoot abduction [20]. A single leg heel rise test is important to evaluate for PTT weakness (Fig. 45.5). Cavovarus feet have characteristic clinical findings. The heel is examined from behind for inversion or from in front for the peek-a-boo sign. The arch height is examined in a weight-bearing position. To evaluate for a flexible or rigid hindfoot, a Coleman block test is performed (Fig. 45.6) [21]. If the hindfoot returns to neutral or slight valgus, then the hindfoot is flexible. First ray and lateral border callosities can be present.

FIGURE 45.5 Too many toes sign. From Masaragian HJ, Massetti S, Perin F, Coria H, Cicarella S, Mizdraji L, Rega L. Flatfoot Deformity Due to Isolated Spring Ligament Injury. J Foot and Ankle Surg. 2020;59:469 478.

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FIGURE 45.6 (A) Shows the peak-a-boo sign when evaluated from the front. (B) Depicts hindfoot varus, greater on the right side. Coleman block testing of left (C) and right (D) foot demonstrating correction of varus deformity (flexible). Courtesy Arthur Manoli II, MD.

45.5

Imaging

It is important to obtain weightbearing radiographs and, if needed, bilateral studies to compare the contralateral hindfoot. Single leg weightbearing films have not been shown to be more reliable even though varus and valgus malalignment are most pronounced during midstance [22]. Stress radiographs generally are not recommended, but intraoperatively, these can be useful to distinguish ankle instability from subtalar instability. The lateral projection can help define the degree of longitudinal arch collapse. The lateral talometatarsal angle (Meary’s angle) (Fig. 45.7), talocalcaneal angle, calcaneal pitch angle, medial cuneiform to fifth metatarsal height, medial arch angle, and talar angle can all be measured on this view [13]. Lateral talometatarsal angle may be the best overall as it assesses the degree of plantigrade tilt of the talar head which is thought to be one of the earliest findings in acquired flatfeet once there is injury to the PTT [13]. Lin and colleagues stated that if the calcaneal pitch angle and the lateral talometatarsal angle are normal, then a PTT tear is unlikely to be present [23]. In addition, the lateral radiograph can evaluate an open vs a closed sinus tarsi and check for anterior beak calcaneal fractures [1]. The oblique view shows tarsal coalitions, especially the calcaneonavicular type. Many specialized views exist to assist in further diagnosing hindfoot pathology. Examples include the alignment view, dorsoplantar view, Broden’s view, hindfoot alignment view (Fig. 45.8), and long axial hindfoot view (Fig. 45.9). CT scans can be utilized for further evaluation and should be weightbearing or simulated weightbearing, if possible. A CT scan can assess articular configuration of the ankle, subtalar, and talonavicular joints. Coalitions can also be seen. Osteoarthritic changes can be apparent. Other bony abnormalities, such as low-grade lateral process fractures of the talus can be seen that are often not evident on plain radiographs [1]. MRI can determine activity of degenerative changes, such as edema. MRI assesses surrounding soft-tissues and synovial inflammation/impingement [15]. It is useful for the assessment of the PTT/spring ligament and of osteochondral lesions [15]. Ultrasonography (US) is an excellent tool for evaluation of soft tissue, especially the PTT. It can show increased tendon thickness, heterogeneous hypoechoic echotexture, and fibrillation. Hypervascularity on color Doppler is also a

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FIGURE 45.7 Lateral weightbearing radiograph of a patient with flatfoot. The two lines drawn through the talus and the axis of the first metatarsal show the lateral talometatarsal (Meary’s) angle. In a normal foot, these lines are parallel. This article is distributed under the terms of the Creative Commons Attribution-Non Commercial 4.0 International (CC BY-NC 4.0) license (https://creativecommons.org/licenses/by-nc/4.0/) which permits non-commercial use, reproduction and distribution of the work without further permission provided the original work is attributed.

FIGURE 45.8 (A) Hindfoot alignment view with the beam angled 20 degrees caudally. (B) Example of hindfoot alignment view radiograph. This article is distributed under the terms of the Creative Commons Attribution-Non Commercial 4.0 International (CC BY-NC 4.0) license (https://creativecommons.org/licenses/by-nc/4.0/) which permits non-commercial use, reproduction and distribution of the work without further permission provided the original work is attributed.

FIGURE 45.9 (A) Long axial view with the beam angled 45 degrees caudally. (B) Example of long axial view radiograph. This article is distributed under the terms of the Creative Commons Attribution-Non Commercial 4.0 International (CC BY-NC 4.0) license (https://creativecommons.org/licenses/by-nc/ 4.0/) which permits non-commercial use, reproduction and distribution of the work without further permission provided the original work is attributed.

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sign of a damaged PTT. US is a unique modality as it can provide dynamic assessment of tendons. US can also evaluate the spring ligament. It can show thickening and can assess for tears, mainly of the superomedial component [15].

45.6

Goals in treatment

Hindfoot malalignment, pes planus, or pes cavus, can predispose the individual to several conditions, including ankle sprains, plantar fasciitis, metatarsal stress fractures, Achilles tendinopathy, as well as knee, hip, and low back pain. The surgeon must identify patients with arch abnormalities. The main goal is to restore normal biomechanics of the foot, which can be achieved through stretching/strengthening exercises, orthotics, and surgery. Other primary goals are to maintain talonavicular and talocalcaneal motion, to maintain medial and lateral column length, stability, and alignment, and to maintain posterior tibial muscle strength. Secondary goals include maintaining cuboid and fourth/fifth metatarsal and calcaneocuboid motion. Also, functional aftercare, avoidance of complications, and rapid healing are crucial. Surgery should address both osseous and soft tissue pathology [13]. It is important to avoid in situ fusions as this does not correct deformity. It is important to pay attention to soft tissue constraints, and it is crucial to carefully prepare fusion surfaces. No direct correlation exists on presurgical evaluation between the degree of deformity and the degree of pain and disability [10].

45.6.1 Treatment goals in flatfoot/cavus syndromes In flatfoot and cavus syndromes, to promote healing and to restore flexibility and strength, the surgeon must reduce mechanical stress on the injured area [10]. It is important for the surgeon to define the mechanical effect needed to change a pathologic condition. This can be achieved by modifying moment arms around the STJ. Medial displacement of the calcaneal tuberosity can reduce the external eversion force. Lateral displacement increases external eversion force [10]. Other main goals in these syndromes include identifying mechanical impairment and knowing the initial problem and consequences. The goals of treatment in flatfoot are to counter the destructive forces, to correct the structural deformity, and to preserve subtalar motion. Two types of reconstructions must be considered: bony and soft tissue. Bone reconstructive options consist of motion-sparing and motion-sacrificing procedures. Motion-sacrificing procedures include talonavicular arthrodesis, triple arthrodesis, and two-column fusion. These options restore joint alignment at the expense of the function of the STJ. Motion-sparing procedures include lateral column lengthening, tuber realignment osteotomy, and medial column stabilization (selected tarsometatarsal/intertarsal) fusion. Soft-tissue reconstructive options include transfer of the flexor digitorum to the PTT/underside of the medial column and reconstruction of the spring ligament. However, reconstruction of the soft tissues does nothing to correct the deformity or the abnormal joint alignment. The best option in this case is to use bony reconstruction to restore anatomy plus soft tissue reconstruction to balance the dynamic forces on the bone. Equinus contracture is the major deforming force that overloads and causes failure of the PTT [11]. In situ fusions should be avoided. The correct position of arthrodesis is the hindfoot placed in 5 degrees of valgus, the transverse tarsal joint in 0 5 degrees of abduction, and the forefoot in less than 10 degrees of varus. The subtalar, talonavicular, and calcanealcuboid joints can be realigned by pushing the calcaneus away from the talus, opening up the sinus tarsi, and taking care not to tip the calcaneus into varus (Fig. 45.10) [1,7]. Fusion of the STJ leads to loss of range of motion and loss of important subtalar functions, such as adaptation and shock absorption. Fusion of the STJ can eliminate pain from arthritis and provide stability. The surgeon should try to preserve the STJ if no pain exists on the physical exam. If there is no passive range of motion or if the STJ is painful in non-weightbearing conditions, then it should be fused. If midfoot malalignment is present after fusion, then it is usually included. If the STJ is left unfused during ankle fusion, then it will be blocked in varus resulting in decreased dorsiflexion during gait [10]. Maceira and Monteagudo prefer to also fuse the STJ when performing an ankle arthrodesis [10]. They make sure to fuse the STJ in an everted position [10]. If varus malalignment is not corrected, this stress on the lateral ligament repairs or reconstructions can lead to failure [15]. In nonparalytic pes cavus (spasmodic, irritative) the foot is short, thick, muscular, contracted, and full of energy. It is usually painful. The goal of treatment is to eliminate deformity, stop muscle spasms, and to protect against future discomfort/insult. Paralytic pes cavus is historically due to poliomyelitis. The treatment was typically unsatisfactory. The patients had atrophic musculature and lessened reflexes with no active contractions. Passive contractions were present due to disuse. In these patients, pain is usually absent. Treatment should first exhaust all nonoperative management. The aim is to support the anterior foot. Shortening of extensor tendons has not been helpful. The insertion of intraarticular silk strands (method of Bartow and Plummer) tends to limit the movements, which should be avoided [24]. It is often necessary to perform osteotomies to correct the foot in plantigrade position [5].

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FIGURE 45.10 This figure depicts the calcaneus being distracted distally to open up the sinus tarsi. Courtesy Hansen ST. Functional reconstruction of the foot and ankle. Lippincott Williams & Wilkins; 2000.

45.6.2 Treatment goals in rheumatoid arthritis The treatment goal of rheumatoid arthritis is to help restore the longitudinal arch [16]. Outcomes tend to be better with treatments that relieve the strain on the PTT and those that maintain motion [16]. Corrective options all attempt to rectify the relationship between the medial and lateral columns and might include: medial displacement calcaneal

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osteotomy, intracalcaneal lengthening osteotomy (i.e., Evans), medial cuneiform dorsal opening wedge osteotomy (i.e., Cotton), subtalar arthroereisis, arthrodesis including the naviculocuneiform/first tarsometatarsal joint, and calcaneocuboid distraction arthrodesis. Patients with advanced disease, those with limited range of motion, arthritis, and moderate joint tenderness, require more motion-limiting procedures, such as a subtalar, talonavicular, or triple arthrodesis [16]. Isolated talonavicular arthrodesis is recommended for flexible rheumatoid flatfoot to prevent progression of the hindfoot and forefoot deformities, although other studies suggest including calcaneocuboid and/or subtalar fusion, because talonavicular arthrodesis effectively causes the loss of 90% or more of complex hindfoot motion, and leaving the diseased joints untreated can lead to further pain in the area [16]. This can increase stability which subsequently may decrease the risk of nonunion. This procedure can also reduce the risk of future surgery for progressive arthritis. The negative aspects of fusing additional joints include the increased morbidity associated with longer procedure time, more incisions, and more joints that can potentially develop a nonunion [16].

45.6.3 Treatment goals in osteoarthritis For patients with osteoarthritis of the tibiotalar joint, nonsurgical treatment should be attempted initially. Pain medication, shoe modifications, bracing, steroid injections, and activity modification are all reasonable choices. If all these measures do not provide adequate relief, then surgery is indicated. Similar to flatfoot, cavus foot, and rheumatoid arthritis, treatment goals are designed to help decrease pain symptoms, correct alignment and stability, and to maximize postoperative function.

45.6.4 Treatment goals in calcaneal fractures The corrective goals of calcaneal fractures are to restore the joint surface and to restore the calcaneal height, length, and heel width [18]. Step-offs of the posterior facet greater than 1 2 mm of length lead to inferior clinical results [18]. Sangeorzan et al. looked at calcaneus fractures through the posterior facet in cadavers. They showed that the contact area of the posterior facet increased with displacement of 5 mm or more. They concluded that fractures should be fixed with displacement of 2 mm or more [25]. Anterior impingement is caused if the posterior talus is impacted into the calcaneal body. A decrease in calcaneal inclination angle leads to the talus becoming more parallel to the ground. This changes the alignment of the tibiotalar joint as well leading to anterior impingement [26].

45.6.5 Treatment goals in talar fractures and dislocations Due to the tenuous blood supply mentioned early, it is important for urgent reduction of displaced talar neck fractures. Rapid revascularization across the fracture surfaces can help prevent avascular necrosis. When attempting reduction, it is crucial to avoid rotational or varus malalignment. Early motion is required for the prevention of stiffness. Sangeorzan et al. studied cadaveric specimens simulating talar neck misalignment [27]. The results showed no change in the posterior facet contact area or mean high-pressure zone area. It also demonstrated no change in the anterior or medial facet contact area. These authors concluded that 2 mm changes in the talar neck alignment lead to offloading of the anteromedial facets without increasing the pressure to the posterior facet. The shift in load may be transferred to the talonavicular joint or may cause impingement of the sinus tarsi with subsequent pain and degenerative arthrosis [27].

45.7

Corrective options

45.7.1 Distraction subtalar fusion In an isolated subtalar fusion, no change in motion occurs in the calcaneocuboid joint. The talonavicular joint motion is reduced by 56% in dorsiflexion and plantarflexion and 70% in eversion and inversion [2]. During the surgical approach, certain techniques are critical to assure appropriate alignment. For example, the use of a pin distractor helps to control the varus/valgus position of the heel while distracting the STJ. In the sagittal plane, the distractor is applied on the posterior aspect of the STJ. In the coronal plane, manual forces are applied to the heel and/ or an additional pin distractor or foot spreader is used on the medial side. For the transverse plane, the talar head is pushed medially with the aid of a Hohmann retractor inserted into the sinus tarsi. Distraction of the posterior STJ and insertion of a graft .10 mm is difficult to achieve through a lateral standard approach (Fig. 45.11) [28]. A

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FIGURE 45.11 (A) Autograft or allograft is made to fit the size of the gap. (B) The subtalar joint is distracted to fit the graft. (C) The graft is impacted between the posterior talus and the calcaneus. (D) Lateral radiograph after graft placement.

posterolateral approach is used for distraction arthrodesis with an angular correction .15 degrees [28]. Not positioning the talus appropriately before insertion of the graft will result in malposition of the STJ. In most instances, the talus is rotated too externally and/or too posteriorly, resulting in a malunion. It is essential to preserve the interosseous ligament to maintain peritalar stability. Two parallel K-wires are inserted from the heel into the posterior aspect of the talar body. Fully-threaded nonlag screws are inserted under fluoroscopic guidance. Use of screws ,6.5 mm may result in failure of the implants [28] (Fig. 45.12). The goal of surgery is to plantarflex the talus and restore the talocalcaneal angle. If the talus is not plantarflexed (after settling in dorsiflexion as a result of calcaneus malunion and flattening of the calcaneal (Bo¨hlers) angle) it will impinge on the distal anterior tibia at the front of the ankle.

45.7.2 Lateral column lengthening Lengthening of the lateral column is accomplished by two methods, by calcaneal lengthening or by arthrodesis of a distracted calcaneocuboid joint. Lateral column lengthening forces the cuboid medially while also correcting peritalar subluxation. The arch is restored and the long ligaments of the foot are re-tensioned. Indications for calcaneal lengthening are PTT insufficiency and medial ankle instability [29]. The hindfoot must be flexible, and arthritis should be absent from the subtalar, talonavicular, and calcaneocuboid joints. An osteotomy is made on the lateral calcaneus, and it is distracted until the forefoot adductus and the MLA is restored [29]. Graft of appropriate size is filled in this gap. Internal fixation is usually not necessary due to internal compression, but a screw may be used to prevent plantar displacement of bone graft [29,30]. Sangeorzan et al. reported that radiographic changes after lateral column lengthening with medial column stabilization included an increase in the calcaneal length of 4 mm, a decrease in the talonavicular angle of 26.8 degrees, an increase of 12 degrees lateral and 16 degrees anteroposterior of the talometatarsal angle, while the calcaneal pitch increased 10.8 degrees, and the talocalcaneal angle decreased 6.5 degrees [31].

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FIGURE 45.12 Postoperative AP and L radiographs after fusion of the talonavicular and subtalar joints.

Arthrodesis of the calcaneocuboid joint stabilizes the lateral column, and it eliminates extreme hindfoot eversion. The support stress is transferred to the bone and not the ligaments [11]. Foot position affects motion after calcaneocuboid fusion in numerous ways. Calcaneocuboid fusion with the foot in neutral demonstrated no change in the talocalcaneal or the talonavicular joint motion compared to the intact foot [11]. When the calcaneocuboid joint was fused at either extreme range of motion, significant decreases in both talocalcaneal and talonavicular motion were noted [11]. In addition, the effect of calcaneal lengthening had no significant effects on any recorded joint motions [11].

45.7.3 Double hindfoot fusion The indications for a double hindfoot arthrodesis (i.e., fusion of the calcaneocuboid and talonavicular joints) include a severe forefoot abduction with a supple transverse tarsal joint, a flexible STJ, and a forefoot that can be brought back

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FIGURE 45.13 Dorsal plantar radiographs after talonavicular and calcaneocuboid arthrodesis.

into plantigrade position [5]. Simultaneous calcaneocuboid and talonavicular fusions have a few advantages. These two joints fuse much easier than the other joints of the hindfoot [32]. Compared to an isolated talonavicular joint fusion, it is less likely to pull the foot into varus [32]. The double hindfoot fusion also allows for some hindfoot, and talonavicular and calcaneocuboid correction. The talonavicular and calcaneocuboid joints are approached through a medial and a lateral incision, respectively. The articular surfaces are prepared, and fixation is obtained without the use of bone graft (Fig. 45.13) [32]. This technique cannot obtain as much hindfoot correction as the triple arthrodesis [32]. The unfused STJ leads to concern of continued pain if degeneration in that joint is present. Another disadvantage of the double hindfoot fusion is that it goes against the current trend to attempt to spare the talonavicular and calcaneocuboid joints while fusing the STJs in that the foot may have compensatory motion through these joints [32]. If the talonavicular and calcaneocuboid joints are preserved, up to 20 degrees of hindfoot dorsiflexion/plantarflexion motion may remain [16]. Triple arthrodesis is used for fixed deformities [5]. In younger adults, surgeons should consider corrective osteotomies and soft-tissue releases if possible [5].

45.7.4 Subtalar and talonavicular fusion For a rigid hindfoot, a triple arthrodesis through a medial and lateral incision is usually indicated. However, when a severe planovalgus foot deformity is present, stress on the lateral soft tissues after realignment is a concern. Omitting arthrodesis of the calcaneocuboid joint can provide adequate correction while also limiting the complications associated

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FIGURE 45.14 Postoperative radiographs showing placement of fully-threaded screws across the STJ.

with a second incision such as wound dehiscence, symptomatic hardware, and nonunion [4]. Decreased surgical times are reported with this single approach without affecting correction of the deformity [33]. The medial incision is made posterior and inferior to the medial malleolus and proceeds distally to the navicular. All three facets of the talus are exposed. The STJ is aligned in 5 7 degrees of valgus, and two 6.5 mm cannulated screws are used for fixation [4,33,34]. The talonavicular joint is reduced to the anatomic position by aligning the talus with the first metatarsal (in both planes). Two or three 5.0 mm cannulated screws are used to obtain fixation across the talonavicular joint [33] (Fig. 45.14). At a mean follow up of 20.6 months, Philippot et al. reported a mean valgus deformity of 11 degrees and a lateral talometatarsal angle of -7.6 degrees [34]. Good pain relief and deformity correction was obtained in their study [34]. Long term outcomes have not yet been reported, but good short-term functional outcomes are apparent [4,33].

45.7.5 Triple arthrodesis Beyond a subtalar fusion, it is important to evaluate if talonavicular and calcaneocuboid joint fusion should be included. Column inequality, talonavicular degenerative joint disease, and an uncorrectable forefoot deformity, such as inversion, are examples of when a triple arthrodesis is considered [35]. The triple arthrodesis is the gold standard as it provides excellent correction, and it has an excellent fusion rate [35]. However, as it eliminates all complex motion of the hindfoot, this procedure is only good for patients who walk on level ground or for more sedentary patients. During surgery, the surgeon must account for other pathology to obtain a successful outcome. Often, a coexisting equinus contracture will prevent reduction of the hindfoot to the proper position. A gastrocnemius lengthening or Achilles tendon lengthening is needed before the reduction and reconstruction can be carried out [35]. The surgeon must avoid excessive bone resection as this will decrease the STJ height and disrupt the articular relationship of the talonavicular joint. An easy way to align the hindfoot is to place a fingertip into the sinus tarsi. The calcaneus is then pushed forward and distal, with careful attention not to displace the heel into varus or valgus. It is critical to correct not just the varus, but also the internal rotation of the calcaneus underneath the talus when reducing the STJ in the case of a planovalgus foot. As the articular cartilage is denuded, the central area of each joint can be further prepared by making a cavity that is filled with bone graft. This acts as a shear-strain relief that ultimately further aids in fusion of the entire construct. The next step is to align the talocalcaneal joint. The hindfoot should now be in the desired alignment. Fixation across the STJs is achieved by drilling from the heel into the talus, with one screw into the body and the other into the neck. The body screw is placed from the lateral part of the inferior calcaneal tuber into the central talar body. The talar neck screw starts more medial on the tuber (as the neck is medial) and is aimed at the central portion of the talar neck. The STJ is now rigidly fixed in proper alignment. Screws across the talocalcaneal joint can be placed dorsal to plantar or plantar to dorsal. Each technique has its own advantages and disadvantages. Placing the screws dorsal to plantar leads to better fixation due to the ability to use a longer threaded screw, but it has a risk of ankle impingement from a screw too close to the talar head [35]. As the dorsal area of the talar neck is one of the main blood supplies of the talus, a dorsal to plantar trajectory risks damaging the dorsal artery to

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FIGURE 45.15 Images (A) and (B) are preoperative weightbearing radiographs demonstrating severe flatfoot deformity. Images (C), (D), and (E) represent immediate postoperative fluoroscopy of a triple arthrodesis in this patient. Image (D) shows the corrections of the lateral talometatarsal angle and calcaneal pitch.

the talus. Screws in the plantar to dorsal direction can avoid impingement by crossing the STJ into the distal talus. However, these will be smaller threaded lag screws (going from a larger fragment of the calcaneus to a smaller fragment of the talus). Staples are also an option for fixation across the talocalcaneal joint, but the trajectory is not as easy to control compared to screws. Also, compression with staples is not reliable [35]. Regardless of this technique, it is essential to obtain two points of fixation across each joint to prevent rotation or sliding which could lead to a nonunion. Next, the talonavicular joint is addressed. Two screws are placed in a lag fashion across the joint either from the navicular into the talus or in an X pattern across the joint. To lag by technique, the near cortex should be drilled the size of the outer diameter of the screw that is to be inserted. This will prevent any purchase of the screw in the near cortex. The far cortex should be drilled the size of the core diameter. When the screw gains purchase in the far bone, it will lag the far bone toward the near bone. Crossed screws offer the advantage of allowing the screw from the medial talar neck to stabilize the lateral navicular [35]. As the lateral portion of the TN joint is often difficult to fuse, this screw alignment may help reduce nonunion of the TN joint. The calcaneocuboid joint is more of an expansion joint and is not as important for stability in the fusion construct [35]. Often, one screw will suffice. It can be placed from the distal calcaneus process into the cuboid. It is common to find a gap at the calcaneocuboid joint that will require bone grafting. This is in essence performing a lateral column lengthening. Final screw placement can be observed via radiographs (Fig. 45.15).

45.7.6 Pantalar arthrodesis Pantalar arthrodesis is indicated for pes plano abducto valgus and ankle tilt where total ankle replacement is not possible. This arthrodesis includes the tibiotalar, subtalar, talonavicular, and calcaneocuboid joints. Essentially, this is a tibiotalocalcaneal fusion combined with talonavicular and calcaneocuboid fusions. Please see prior techniques for reference.

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It is difficult to obtain fusion, especially in obese patients. It achieves good correction, but creates a “living” prosthesis [7]. A below the knee amputation might be a better choice.

45.7.7 General complications Excessive hindfoot valgus increases stress on the deltoid ligament by 76% [15]. This also leads to lateral subfibular impingement. Residual hindfoot varus causes overloading of the lateral column and pain in the cuboid or in the base of the fifth metatarsal. This may lead to lateral ankle instability and secondary ankle arthritis. Failure to address an Achilles tendon contracture stresses the midfoot and subsequent avascular necrosis, or nonunion can develop. Nonunion most commonly occurs at the talonavicular joint. In addition, secondary arthrosis can ensue [5].

45.7.8 Postoperative management requirements in hindfoot fusions While surgeons’ preferences may differ, typically after postoperative arthrodesis, a patient is required to be nonweightbearing and the foot and ankle are placed in a splint for 10 14 days to allow the soft tissues to heal. At this time, the sutures are removed. The patient continues non-weightbearing restrictions for another 4 weeks. The patient is evaluated for healing at 6 weeks, and if adequate, may start progressive weightbearing. Owing to the loss of accommodation of the hindfoot after fusion, a substitute must be made to account for the absence of shock absorption. This can include proper shoes, soft heel activities (low-impact), and weight control.

45.8

Areas of future interest

As more research is performed, attention should be paid to the use of cartilage resurfacing to improve pain and function while avoiding a fusion procedure. Also, the use of biologics should be considered for the treatment of hindfoot pathology. This may have a role in symptomatic treatment that can delay surgery. Supplementation with biologics may lead to increased healing rates and may also decrease the nonunion rates. For cases of fracture and deformity, the role of the contralateral limb for evaluation and imaging should be explored. Diagnosis of hindfoot deformity of the contralateral limb can provide important information for planning of the surgical foot. Correction of presumed malalignment during ORIF could help in preventing reinjury as well as failure of the repair.

References [1] Hansen ST. Functional reconstruction of the foot and ankle, 1. Philadelphia: Lippincott Williams & Wilkins; 2000. p. 17 32. [2] Savory KM, Wulker N, Stukenborg C, Alfke D. Biomechanics of the hindfoot joints in response to degenerative hindfoot arthrodeses. Clin Biomech 1998;13(1):62 70. [3] Ryerson EW. The classic: arthrodesing operations on the feet. Clin Orthop Relat Res 1977;122:4 9. [4] Sammarco VJ, Magur EG, Sammarco GJ, Bagwe MR. Arthrodesis of the subtalar and talonavicular joints for correction of symptomatic hindfoot malalignment. Foot Ankle Int 2006;27(9):661 6. [5] Wapner KL. Triple arthrodesis in adults. J Am Acad Orthop Surg 1998;6:188 96. [6] Bowers CA, Catanzariti AR, Mendicino RW. Traditional ankle arthrodesis for the treatment of ankle arthritis. Clin Podiatr Med Surg 2009;26 (2):259 71. [7] Takahashi KZ, Worster K, Bruening DA. Energy neutral: the human foot and ankle subsections combine to produce near zero net mechanical work during walking. Sci Rep 2017;7(1):15404. [8] Erdemir A, Hamel AJ, Fauth AR, Piazza SJ, Sharkey NA. Dynamic loading of the plantar aponeurosis in walking. J Bone Jt Surg Am 2004;86 (3):546 52. [9] Krahenbul N, Horn-Lang T, Hintermann B, Knupp M. The subtalar joint: a complex mechanism. EFORT Open Rev 2017;2(7):309 16. [10] Maceira E, Monteagudo M. Subtalar anatomy and mechanics. Foot Ankle Clin 2015;20(2):195 221. [11] Sands A, Early J, Harrington RM, Tencer AF, Ching RP, Sangeorzan BJ. Effect of variations in calcaneocuboid fusion technique on kinematics of the normal hindfoot. Foot Ankle Int 1998;19(1):19 25. [12] Astion DJ, Deland JT, Otis JC, Kenneally S. Motion of the hindfoot after simulated arthrodesis. J Bone Jt Surg Am 1997;79(2):241 6. [13] Lin YC, Kwon JY, Ghorbanhoseini M, Wu JS. The hindfoot arch: what role does the imager play? Radiol Clin North Am 2016;54(5):951 68. [14] Ramachandra P, Kumar P, Kamath A, Maiya AG. Do structural changes of the foot influence plantar pressure patterns during various stages of pregnancy and postpartum? Foot Ankle Spec 2017;10(6):513 19. [15] Krause F, Seidel A. Malalignment and lateral ankle instability: causes of failure from the varus tibia to the cavovarus foot. Foot Ankle Clin 2018;23(4):593 603.

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[16] Aronow MS, Hakim-Zargar M. Management of hindfoot disease in rheumatoid arthritis. Foot Ankle Clin 2007;12(3):455 74. [17] Snedeker JG, Wirth SH, Espinosa N. Biomechanics of the normal and arthritic ankle joint. Foot Ankle Clin 2012;17(4):517 28. [18] Xu C, Li M, Wang C, Li K, Liu H. Ankle-hindfoot after calcaneal fractures; a biomechanical study. Orthop Traumatol Surg Res 2017;103 (5):709 16. [19] Weinstein SL, Flynn JM. The foot. Lovell and winter’s pediatric orthopaedics, 29. Philadelphia: Lippincott Williams & Wilkins; 2006. p. 1390 1. [20] Beals TC, Pomeroy GC, Manoli 2nd A. Posterior tibial tendon insufficiency: diagnosis and treatment. J Am Acad Orthop Surg 1999;7 (2):112 18. [21] Manoli 2nd A, Graham B. Clinical and new aspects of the subtle cavus foot: a review of an additional twelve year experience. Fuß Sprunggelenk 2018;16. Available from: https://doi.org/10.1016/j.fuspru.2017.11.006. [22] Reilingh ML, Beimers L, Tuijthof GJ, Stufkens SA, Maas M, van Dijk CN. Measuring hindfoot alignment radiographically: the long axial view is more reliable than the hindfoot alignment view. Skelet Radiol 2010;39(11):1103 8. [23] Lin YC, Mhuircheartaigh JN, Lamb J, et al. Imaging of adult flatfoot: correlation of radiographic measurements with MRI. AJR Am J Roentgenol 2015;204(2):354 9. [24] Davis GG. The treatment of hollow foot (Pes Cavus). J Bone Jt Surg Am 1913;s2 11(2):231 42. [25] Sangeorzan BJ, Ananthakrishnan D, Tencer AF. Contact characteristics of the subtalar joint after a simulated calcaneus fracture. J Orthop Trauma 1995;9(3):251 8. [26] DiDomenico LA, Butto DN. Subtalar joint arthrodesis for elective and posttraumatic foot and ankle deformities. Clin Podiatr Med Surg 2017;34(3):327 38. [27] Sangeorzan BJ, Wagner UA, Harrington RM, Tencer AF. Contact characteristics of the subtalar joint: the effect of talar neck misalignment. J Orthop Res 1992;10(4):544 51. [28] Hinterman B, Ruiz R. Distraction subtalar fusion. 2nd ed. Foot and ankle surgery, 43. Philadelphia: Elsevier; 2018. p. 361 70. [29] Hinterman B, Valderrabano V. Lateral calcaneal lengthening osteotomy for supple adult flatfoot. 2nd ed Foot and ankle surgery, 33. Philadelphia: Elsevier; 2018. p. 264 9. [30] Sands AK, Tansey JP. Lateral column lengthening. Foot Ankle Clin 2007;12(2):301 8 vi vii. [31] Sangeorzan BJ, Mosca V, Hansen Jr. ST. Effect of calcaneal lengthening on relationships among the hindfoot, midfoot, and forefoot. Foot Ankle 1993;14(3):136 41. [32] Clain MR, Baxter DE. Simultaneous calcaneocuboid and talonavicular fusion. Long-term follow-up study. J Bone Jt Surg Br 1994;76 (1):133 6. [33] Lee MS. Medial approach to the severe valgus foot. Clin Podiatr Med Surg 2007;24(4):735 44 ix. [34] Philippot R, Wegrzyn J, Besse JL. Arthrodesis of the subtalar and talonavicular joints through a medial surgical approach: a series of 15 cases. Arch Orthop Trauma Surg 2010;130(5):599 603. [35] Breceda AP, Sands AK. Triple arthrodesis. 2nd ed. Foot and ankle surgery, 44. Philadelphia: Elsevier; 2018. p. 371 7.

Chapter 46

Biomechanics of Foot and Ankle Fixation Justin K. Greisberg and Christina E. Freibott Department of Orthopedic Surgery, Columbia University Medical Center, New York, NY, United States

Abstract Successful fixation of fractures or fusions of the foot and ankle require rigid internal fixation, preservation of the biology of the bone, and protection of the bone from repetitive stress during the healing process. This can be achieved through a variety of methodologies. Screws can be placed across a fracture or fusion, during which overdrilling the proximal fragment allows compression of the bone fragments, which promotes healing. A variety of plating constructs can be utilized to achieve compression, neutralization, buttressing, or bridging. A post and screw construct can be implemented to improve compression and mechanical fixation, thus increasing fusion rates. Intramedullary nails are flexible but allow secondary healing of the bone with callus. Lastly, long, thick screws down the center axis of a fused segment can be utilized to form an internal scaffold, or “beaming,” and subsequently optimize the biomechanics of fixation.

46.1

Introduction

During surgery for fracture repair, osteotomy, or arthrodesis (fusion), internal fixation techniques are used to hold bone fragments in position while the bone is healing [1,2]. In the foot and ankle, weight-bearing forces across the bones (prior to healing) will stress the internal fixation construct. These forces can be several times body weight. With sufficient repetitive stress, components of the internal fixation will fatigue and subsequently fail. Therefore, if a screw breaks after being used to repair a fracture in the foot, it is unlikely to be due to a single large stress, but rather the result of repetitive micromotion and fatigue [2 4]. With any internal fixation, there is a “race” between successful healing of the bone and failure of the fixation. For successful long-term outcomes after fixation surgery, the construct must be strong enough to endure any micromotion until osseous union occurs [5]. Nonweight bearing through the period of healing can be beneficial. In the foot and ankle, primary bone healing is routinely required for fractures, fusions, and osteotomies. With primary bone healing, osteoblasts lay down osteoid across the fracture. These osteoblasts cannot “jump” across a gap, so direct bone healing requires the bone fragments to be tightly apposed to each other. In the case of fractures treated nonoperatively, bone healing tends to be indirect, or secondary. In this case, the initial trauma leads to hematoma formation along with inflammation. Over time, some micromotion results in callus formation within the hematoma. This callus progresses from fibrous tissue to fibrocartilage and eventually to woven bone [6]. Further remodeling organizes the woven bone into cortical bone according to the mechanical environment. Radiographs will often show a “cloud” of callus around the fragments (Fig. 46.1). Micromotion is considered desirable for nonsurgical fracture healing. If there is absolutely no micromotion, as might occur if bone fragments are fixed with a rigid plate, then the secondary healing, and callus formation, will not occur. In fact, the presence of fracture callus on radiographs following rigid fixation may be a sign of failure of fixation (because motion is present), as opposed to successful healing. In cases of rigid fixation, bone healing will be direct, or primary, requiring bone fragments to be tightly apposed. This is a key point; if rigid internal fixation is being used, then there cannot be any gap between the bone fragments. Therefore, surgeons use internal fixation that promotes compression of the bones. Compression is also important in improving the strength of the internal fixation construct. Interdigitation of the bone fragments increases friction at the Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00008-1 © 2023 Elsevier Inc. All rights reserved.

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FIGURE 46.1 Lateral view of a tibia fracture several months after nailing with the fracture callus is visible (indicated by the arrows).

fracture or osteotomy, and makes the construct more resistant to shearing or motion. Therefore, compression fixation can help to protect the internal fixation from fatigue [7,8]. However, the surgeon is not simply a carpenter, screwing the bones together. Although it is essential to achieve sufficiently durable fixation, it is just as important to preserve the vitality of the bone. The osteoblasts require an intact blood supply to the bone. If aggressive surgical exposure devascularizes the bone, the osteoblasts will not be able to grow bone across to the other fragment. In this sense, the surgeon is more of a gardener than a carpenter, providing the right environment for the bone to grow. If excessive periosteal stripping is performed to place internal fixation, no healing will occur. To summarize, successful internal fixation with primary bone healing requires: 1. rigid internal fixation with compression to resist gapping 2. protection of the bone from repetitive stresses during healing 3. preservation of the vasculature of the bone. It is noteworthy that internal fixation can occasionally be used as a splint, rather than rigid fixation. This is especially true with intramedullary nails (discussed later in this chapter), where union may occur through secondary bone healing. In these cases, it is still essential to preserve the vasculature of the bone and limit repetitive stress to within the capabilities of the nail, but fixation with a gap may be acceptable. This chapter discusses the hardware and techniques used to achieve internal fixation, with particular reference to procedures commonly used for the foot and ankle.

46.2

Screws

Bone screws used for internal fixation are similar in principle to other, commonly used screws, for example in woodworking. The key properties of a typical screw are shown are related the threads and the diameter (Fig. 46.2).

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FIGURE 46.2 A typical 6.5 mm bone screw with key dimensions identified.

When a screw is tightened onto a plate or into bone, the head of the screw puts pressure on the plate or bone where it contacts, and the threads of the screw bind against the head side of the helical thread grove they have cut into the bone—these opposing reactions compress the material between the head and the far end of the screw tip. A larger diameter head provides more surface area for those forces and decreases the chance of the screw burrowing through the bone or plate. At the same time, a larger head is more prominent in the soft tissues, so the right balance must be achieved. The core (root) diameter of the screw is most important for the strength of the screw in bending and fatigue. This strength is proportional to the radius of the core, to the fourth power. A small increase in core diameter can lead to a relatively large increase in screw strength (and thus resistance to fracture). That is why many surgeons fix the midfoot in a case of a metatarsal being displaced from the tarsus (also known as Lisfranc injury) with 4.0 mm screws. These screws have a core diameter of 2.9 mm, as opposed to the 2.5 mm core diameter for the 3.5 mm screws that are commonly used for small fragment fixation. This small increase in diameter results in an increase of about 80% in strength [9]. The difference between the outer diameter and the inner diameter of the screw is the depth of the threads. The pullout strength of a screw, and the ultimate tightening force of a screw, is related to this thread depth. Pullout strength is also related to the “density” of the threads, or the number of threads per inch. However, if the threads are too dense, leaving too little bone between each thread, then the interface with the bone is compromised, decreasing the pullout strength. Traditionally, “cortical” screws have had smaller thread depths, for tightening into dense cortical bone. Cancellous screws have thicker thread depths, with less thread density, for looser bone. These differences have become more blurred over time, as screw development has progressed. Cannulated screws have a hollow core, to allow passage over a wire. This makes placement of the screw easier, as the surgeon can initially place a thin wire, and check its position with X-ray, before placing the screw. For any given

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screw outer diameter, the core diameter of a cannulated screw must be larger, to make up for the lost metal from the hollow core to provide equivalent mechanical strength to a comparable solid screw. Cannulated screws also tend to be more expensive than their solid counterparts. For internal fixation of foot fractures or fusions with small fragments, it is common to use screws with outer diameters of 2.0, 2.4, or 2.7 mm. The relative strength of these screws is not great, and screw fracture will occur in some cases, for example with interphalangeal joint fixation of the toes if osseous healing does not occur quickly. Midfoot fusions and ankle fractures are often fixed with 3.5 mm screws. Hindfoot or ankle fusions may be fixed with 6.5 mm screws, but even these will eventually fatigue and fracture in the case of nonunion. When placing a screw across a fracture or fusion, it is routine to overdrill the proximal fragment of bone, so the screw threads can slide unopposed through it. This allows compression of the bone fragments which, as described earlier, promotes primary osseous healing. Alternatively, a partially threaded screw may be used, but a fully-threaded screw placed in compression will generally have better purchase. The partially threaded screw is limited by the small number of threads on the screw, whereas a fully threaded screw will have threads in the entire fragment. If a fully threaded screw is placed without overdrilling the proximal fragment, then the screw is considered to be a position screw. There will be no compression between the fragments. When fixing midfoot injuries, this is a common practice, to avoid compression of the articular surfaces. The surgeon is looking to hold the joint reduced while the capsule heals, but no bone-to-bone healing is intended.

46.3

Plates

Internal fixation plates can be used to achieve compression, neutralization, buttressing, or bridging. These plates come in various shapes, thicknesses, and materials. A thicker plate will be more resistant to bending and fatigue. A thinner plate will be less prominent on the surrounding soft tissues. Plates can be used to achieve compression, neutralization, buttressing, or bridging. The names given to the different types of plate (tubular, compression, or reconstruction) tend to be more historical than descriptive. Plates are not a mode of fixation on their own; it is what a surgeon does with a plate that really matters. Tubular plates are thin, relative to thicker compression plates. Reconstruction plates are thicker but are also malleable, so they can be contoured more easily than compression plates [10]. Neutralization plates provide additional stabilization to a compression construct. First, a lag screw (separate from any plate) is used to compress bone fragments. Then a plate is applied across the fragments for further support. This is commonly performed for fixation of routine lateral malleolus fractures (Fig. 46.3). A compression plate is applied to one bone fragment first with a screw (Fig. 46.4). The screw holes in a compression plate are typically oblong or oval, with an inclined plane at one end. The hole for the first screw in the other bone fragment is drilled off center, so that as the screw is tightened down (and as the head of the screw hits the inclined plane at the end of the hole), the bone fragments are pulled together. This can achieve compression at the fracture/ osteotomy.

FIGURE 46.3 A lag screw is first used to compress the bone fragments together, proximal fragment is overdrilled, then a neutralization plate is applied for additional support.

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FIGURE 46.4 Compression can be achieved through a plate. The plate is fixed to one fragment of bone with a screw or two. Then a screw is placed off-center in a hole of the plate, into the second bone fragment. When this screw is tightened, the head of the screw interacts with the plate pushing the screw and thus the underlying bone into the other bone fragment, to compress across the gap.

FIGURE 46.5 With a partial articular fracture, a buttress plate gives support and squeezes the bone fragments together.

A plate can also be used to buttress a partial articular fracture, such as the tibial plateau or in the distal tibia, by undercontouring the plate so it pushes the articular fragment firmly into position (Fig. 46.5). Some surgeons refer to these as spring plates, since the plate pushes the fragment into the bone. This is commonly used for large medial malleolus fragments in supination and adduction ankle fractures. For fractures comprising multiple bone fragments (comminution), compression is not possible. In these cases a plate can bridge, or span, the fracture (i.e., bridge plating). The portion of the plate passing the comminution may experience substantial bending stress, so the plate needs to be strong enough to support the anticipated stress over time (Fig. 46.6). For any plating construct, failure typically occurs by screws loosening in the bone. The plate will lose its fixation as screws loosen over time [11]. Screws tend to loosen independently of each other, but once the first one loosens, the others begin seeing more stress, so their likelihood of failure increases. The introduction of locking plates gave surgeons an additional option for fixation, although not a new method of plating. Locking screws have threads in the head of the screw, which engage with the corresponding hole on the plate. The screw therefore becomes fixed to the plate, and at a fixed, designed angle. If a plate is attached to a bone with all locking screws, and then stressed over time, the locking screws cannot loosen independently. Failure comes from all screws loosening together, since they are all locked and cannot move (or wiggle loose) independently. Therefore, locked constructs are more resistant to loosening from cyclic loading. Locking fixation may be more durable, especially in weaker bone, and somewhat limits the screw trajectory, but the concepts of plating remain the same [12].

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FIGURE 46.6 For fractures with comminution, a bridge plate spans across the fragments, and healing should occur with callus.

Internal fixators may seem to be a novel concept in plating, but really are just bridge plates. If a locked plate is applied with all locking screws, none of the screws will tighten the plate to the bone. The plate may not touch the bone at all. This can be called an internal fixatory, which is similar to an external fixator in theory. Then internal fixators function much like the bridge plate, as described earlier.

46.4

Post and screw constructs

When placing a lag screw, the ideal screw would pass through the center of the bones, perpendicular to the fusion (or fracture) site. Owing to the close proximity and overlapping nature of the bones involved in ankle or hindfoot fusions, in these cases, it is only possible to place the lag screw from the edge of the bones and joints. In many cases, compression is off center, and incomplete. This is perhaps most problematic in a talonavicular fusion but is present to some degree in all foot and ankle fusions. Improved compression can be achieved by first placing a post in the navicular. A lag screw is then drilled through the post (Fig. 46.7). As the screw locks into the post during final tightening, the post pulls the entire navicular into the talus. Such post and screw constructs provide improved mechanical fixation, and increased fusion rates [13]. Recently, some manufacturers have designed plates that allow a screw to lag across a fusion, through the plate. This provides fixation similar to the post and screw construct. The plate is fixed to one bone of a fusion (e.g., the metatarsal in a metatarsophalangeal fusion), and then a lag screw is placed through the plate from this bone (metatarsal side) to the other bone (phalanx). The lag screw tightens into a special recess in the plate, and as it is tightened, it compresses across the fusion. Finally, screws are placed from the plate into the far bone (phalanx), neutralizing the compression.

46.5

Nails

When a carpenter’s nail is hammered into a piece of wood, it displaces the wood and achieves a press-fit. While it is very important to drill a piece of wood for the core diameter when inserting a screw, if a carpenter did the equivalent predrilling for a nail, the nail would be loose. Intramedullary nails were introduced in the middle of the 20th century for mid-shaft tibia and femur fractures. The nails were essentially a thin sheet of steel rolled into a long tube, and achieved a press fit in the medullary canal, much as a nail does in wood. Bone is not exactly like wood, so some predrilling (reaming) of the bone is necessary to make room for the nail, but still a press-fit inside the bone is desirable. Reaming also stimulates biologic activity of the bone [14 16], which is good for healing, so all nailing is done with some degree of reaming. Nails have holes at either end for the placement of interlock screws. These screws can help prevent rotation of the bone fragments, and also can keep the bone from telescoping, or collapsing, along the nail for comminuted fractures. If the nail is completely loose inside the bone, it can still act as a splint for the bone, and if interlock screws are used, the nail could even be thought of as an intramedullary plate. This is commonly seen with nails used for intertrochanteric hip fracture fixation.

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FIGURE 46.7 In this schematic, a superior view of a left hindfoot, the post is placed in the cuboid. A lag screw through the post, into the calcaneus, will pull the post, and thus the cuboid, toward the calcaneus. Modified from Greisberg J, Vosseller JT, Ferry C, Nash C, Gardner TR. A new method for achieving compression in hindfoot arthrodesis. Int Orthop. 2015;39(11):2267 74. https://doi.org/10.1007/s00264-015-2855-y.

Intramedullary nails are somewhat flexible, and do not achieve absolutely rigid fixation of a fracture [17]. The fracture will thus heal by secondary healing, with callus. The appearance of callus is a good sign when a nail is used to stabilize a tibia fracture, which is the opposite of what one might see with a compression plate. When stabilizing a long bone fracture (femoral or tibial shaft) or a long bone fusion (ankle or knee), there will be a bending force on the internal fixation when forces are applied (such as with weight bearing prior to healing). This bending force is proportional to the distance from the center of the bone (the axis of weight bearing). If internal fixation is down the center of the bone, these bending forces are considerably less than if the fixation is outside the bone. Therefore, intramedullary nails may be biomechanically better than “extra medullary” plates for stabilizing long bones [18]. In the context of foot and ankle fractures, nails have been adapted for tibio-talo-calcaneal fusions. The nails for this application are shorter and inserted from an incision in the sole of the foot. The reaming required for insertion may stimulate healing of the fusions, and the reamings may act as bone graft. The nails are quite stout, and are positioned in the center of the joints, and so are biomechanically strong for this application. Biomechanical testing for stiffness or load to failure in general will show better results for these “ankle fusion nails” when compared to a lateral plate, especially in the presence of bone loss [19]. But these tibio-talo-calcaneal fusions are usually being performed because of severe bone loss or deformity, where strength of the construct may not be the most important factor for achieving successful union. Furthermore, by design, these nails do not achieve a tight press fit in the talus or calcaneus. Interlocking screws placed through those bones augment fixation. The nails may be more of an intramedullary plate in the hindfoot. The nails do achieve a tight fit in the medullary canal of the tibia, but there is a stress riser at the proximal end of the nail, which is usually in the distal third of the tibial shaft. Tibial fracture at the proximal tip of the nail occurs intermittently because of this stress. Use of a full-length nail that extends to the proximal tibial metaphysis will eliminate that risk.

46.6

Beams

Charcot neuroarthopathy of the hindfoot may result in dislocation at the talonavicular or transverse tarsal joints. All stability of the hindfoot and midfoot is often lost. When realigning and stabilizing such a foot, it is important to achieve strong fixation. Healing is generally slow, and the end result is often a fibrous union, not fusion. Furthermore, because of the neuropathy, these patients are very prone to early weight bearing. Their balance is poor, and they will not realize they are bearing weight on the foot at all. In charcot cases, the surgeon is faced with slow healing and early weight bearing, so the race of successful healing versus fixation failure is tipped the wrong way. Traditional plates and/or screws for hindfoot fusion in Charcot cases will usually fail. Stronger fixation is required. Placing fixation down the center of the axis of the fused segment will optimize the biomechanics of fixation, as will using thick implants (lots of metal to resist fatigue). In the case of a tibio-talo-calcaneal fusion, an intramedullary ankle fusion nail can achieve this. For the hindfoot or midfoot, long, thick screws can be used in a similar way (Fig. 46.8). Stress shielding does not seem to be an issue with any internal fixation in the foot/ankle.

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FIGURE 46.8 Long, thick screws are used as internal beams, or rebar, to support the columns of the foot. This patient had Charcot neuropathic arthropathy of the midfoot.

The concept of internal reinforcing beams was popularized in the 1990s by Hansen et al. by drawing parallels to rebar in concrete construction [3]. (Technically, rebar in concrete is used to promote strength in tension.) These long screws provide an internal scaffold, within the osseous anatomy. Later, surgeons referred to this fixation as beaming. In a typical hindfoot charcot deformity, the navicular may dissolve, so that the forefoot (metatarsals) are completely dislocated from the hindfoot (talus and calcaneus). In a noncharcot hindfoot fusion, lag screws hold the talus to the navicular, but the joints being fused are stable, not dislocated, so the forces across those screws are not extreme. With a neuropathic dislocation, there is no intrinsic stability at all, and routine lag screws will fail quickly. A more durable construct involves placing a thick solid screw (typically 6.5 mm outer diameter screw with a core diameter of 4.0 mm) through the talus and down the canal of the first metatarsal. It is possible to place a similar screw from the calcaneus down the axis of the fourth or possibly fifth metatarsal. Alternately, the screws can be placed retrograde, from the metatarsal head into the hindfoot bones.

46.7

Areas of future research

Given that noncompliance with nonweight bearing is an ongoing challenge, there is a need for implants that are stronger and have greater fixation strength to the bone. It is still too common to see nonunions, both with screw fracture and also with loss of screw purchase in the bone. This would reduce the time of nonweight bearing. As new technical developments continue to create new implants (post and screw, locking, etc.), it will be important to critically assess whether the increased cost of these implants can be justified by improved outcomes. More rigid fixation may offer higher union rates or may permit earlier weight bearing. Both of these concepts, if examined through a medical, biomechanical, and economic lens, will be valuable additions to the field of biomechanics research in the next 5 10 years.

References [1] Rockwood and Green’s fractures in adults. R2 Digital Library. ,https://www.r2library.com/Resource/Title/1605476773. [accessed 22.12.18]. [2] Browner BD, Jupiter JB, Krettek C, Anderson P, editors. Skeletal trauma: basic science, management, and reconstruction, vol 1. 5, illustrated. Elsevier/Saunders; 2014. [3] Elements of fracture fixation—3rd ed. ,https://www.elsevier.com/books/elements-of-fracture-fixation/thakur/978-81-312-4237-7. [accessed 22.12.18]. [4] Schrenker R. Learning from failure: the teachings of Petroski. Biomed Instrum Technol 2007;41(5):395 8. [5] Laurence M. The elements of fracture fixation. J Bone Jt Surg Br 2008;90-B(7). Available from: https://doi.org/10.1302/0301-620X.90B7.21132 980-980. [6] Recknagel S, Bindl R, Brochhausen C, et al. Systemic inflammation induced by a thoracic trauma alters the cellular composition of the early fracture callus. J Trauma Acute Care Surg 2013;74(2):531 7. Available from: https://doi.org/10.1097/TA.0b013e318278956d. [7] Brunner CF, Weber BG. Internal fixation plates with a specialized form or function. Spec tech intern fixation. Berlin, Heidelberg: Springer Berlin Heidelberg; 1982, p. 145 65. Available from: http://doi.org/10.1007/978-3-662-02345-7_8.

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[8] Brunner CF, Weber BG. Special techniques in internal fixation. Berlin, Heidelberg: Springer Berlin Heidelberg; 1982. Available from: http:// doi.org/10.1007/978-3-662-02345-7. [9] Tencer AF, Johnson KD. Biomechanics in orthopedic trauma: bone fracture and fixation. 1st ed. Lippincott Williams; 1994. [10] Mu¨ller ME, Allgo¨wer M, Schneider R, Willenegger H. Manual of internal fixation. Berlin, Heidelberg: Springer Berlin Heidelberg; 1979. Available from: http://doi.org/10.1007/978-3-642-96505-0. [11] Gautier E, Sommer C. Guidelines for the clinical application of the LCP. Injury. 2003;34(Suppl. 2):B63 76. [12] Kiner DW, Rouhipour V, Kellam JF. Biomechanics of locked plate fixation. Tech Orthop 2007;22(3):151 5. Available from: https://doi.org/ 10.1097/BTO.0b013e318149fc0b. [13] Greisberg J, Vosseller JT, Ferry C, Nash C, Gardner TR. A new method for achieving compression in hindfoot arthrodesis. Int Orthop 2015;39 (11):2267 74. Available from: https://doi.org/10.1007/s00264-015-2855-y. [14] Greksa F, To´th K, Boros M, Szabo´ A. Periosteal microvascular reorganization after tibial reaming and intramedullary nailing in rats. J Orthop Sci 2012;17(4):477 83. Available from: https://doi.org/10.1007/s00776-012-0222-z. [15] Leung KS, Cheung ENM. Biology and physiology of intramedullary reaming in the fixation of fractures. In: Kempf I, Leung KS, Grosse A, Haarman HJTM, Seidel H, Taglang G, editors. Practice of intramedullary locked nails. Berlin, Heidelberg: Springer Berlin Heidelberg; 2002, p. 31 41. Available from: https://doi.org/10.1007/978-3-642-56330-0_4. [16] Rand JA, An KN, Chao EY, Kelly PJ. A comparison of the effect of open intramedullary nailing and compression-plate fixation on fracture-site blood flow and fracture union. J Bone Jt Surg Am 1981;63(3):427 42. [17] Curtiss PH. Practice of intramedullary nailing. Am J Surg 1968;116(5):803. Available from: https://doi.org/10.1016/0002-9610(68)90374-7. [18] Winquist RA, Hansen ST, Clawson DK. Closed intramedullary nailing of femoral fractures. A report of five hundred and twenty cases. J Bone Jt Surg Am 1984;66(4):529 39. Available from: https://doi.org/10.2106/00004623-198466040-00006. [19] Hansen ST. Functional reconstruction of the foot and ankle. illustrated. Lippincott Williams & Wilkins; 2000.

Chapter 47

Ankle Arthroplasty and Ankle Arthrodesis Daniel C. Norvell1,2, Sagar S. Chawla3 and William R. Ledoux1,3,4 1

RR&D Center for Limb Loss and MoBility (CLiMB), Veterans Affairs Puget Sound Health Care System, Seattle, WA, United States, 2Department of

Rehabilitation Medicine, University of Washington, Seattle, WA, United States, 3Department of Orthopaedics and Sports Medicine, University of Washington, Seattle, WA, United States, 4Department of Mechanical Engineering, University of Washington, Seattle, WA, United States

Abstract Ankle arthrodesis and ankle arthroplasty are performed on patients when the pain and functional limitations of end-stage ankle arthritis (ESAA) is no longer responsive to conservative management. The choice between arthrodesis or arthroplasty depends on patient and surgeon preferences as well as patient comorbidities. The goal of the arthrodesis is to fuse all relevant bones so that there is no longer motion between them, which eliminates the pain; however, the patient will never walk normally again. In contrast, arthroplasty provides new joint surfaces and helps in preserving natural ankle motion. While arthrodesis has been the only surgical option and considered as the gold standard for decades, arthroplasty has slowly become a viable alternative with first-generation implants giving way to second-and third-generation designs, which are much safer and more effective than the original implants. When considering ankle arthroplasty, the biomechanics of the ankle must be considered, including limb and ankle alignment, bony and ligamentous ankle anatomy, and ankle motion in the three cardinal planes. Because arthrodesis is inherently motion limiting, the factors related to motion are not relevant; however, there are biomechanical concerns that should be considered related to bony morphology and alignment prior to surgery. There are also numerous biomechanical considerations and/or complications from treating patients with ESAA with an ankle arthrodesis or arthroplasty that should be addressed, including ankle alignment, gait mechanics, arthritis at distal joints, component wear or failure, cadaveric gait simulation, and computational modeling. Biomechanical studies have explored outcomes comparing arthroplasty to arthrodesis in terms of the overall arc of motion, sagittal plain range of motion, sagittal power, step counts, and restoration of gait to as close to normal as possible. Because of the highly technical nature of ankle arthroplasty, studies have demonstrated better outcomes in surgeons with more training and experience. Clinical studies have demonstrated successful outcomes in patients who have undergone arthrodesis; however, drawbacks have included the extended time required to achieve fusion, potential for nonunion, leg length discrepancy, malalignment, stress fractures, continued pain, and arthritis in distal joints requiring secondary procedures. This has led to more research on ankle arthroplasty where ankle motion is retained. Several studies suggest that short-term clinical and functional outcomes appear similar if not superior after arthroplasty, but reoperation rates or major complications are greater; however, more recent comparative effectiveness studies indicate superior patient-reported outcomes and reduced revision rates in arthroplasty compared to arthrodesis. Longer term follow-up comparing these two procedures beyond 10 years is still lacking but will likely emerge in the near future from ongoing studies.

47.1

Introduction

Ankle fusion, that is, ankle arthrodesis, or alternatively ankle joint replacement, that is, ankle arthroplasty, are performed on patients when the pain and functional limitations of end-stage ankle arthritis (ESAA) is no longer responsive to medications, physical therapy, or other conservative measures. The cause of pain and dysfunction comes from degeneration of the cartilage that covers the ends of the three primary bones that make up the ankle joint: the tibia, fibula, and talus. Unlike the hip or the knee, where primary osteoarthritis (OA) is common, most occurrences of ankle OA are due to a previous traumatic insult [1]. Whether a patient receives an arthrodesis or an arthroplasty depends on many factors, including the preferences of the surgeon and the patient, as well as the age, activity-level, and body mass index (BMI) of the patient. Unfortunately, there are no evidence-based clinical guidelines to determine which treatment a Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00009-3 © 2023 Elsevier Inc. All rights reserved.

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particular patient should receive. This chapter includes a description of ankle arthrodesis and arthroplasty techniques, a summary of the biomechanical factors in presurgical assessment, a summary of the biomechanical considerations/complications and biomechanical outcomes of both procedures, as well as a review of clinical outcomes associated with both procedures. Finally, we conclude with future biomechanical research areas that should be explored.

47.2

Brief description and history of surgical techniques

47.2.1 Ankle arthrodesis The goal of the arthrodesis procedure is to fuse the tibia, fibula, and talus so that there is no longer motion between them, which in turn reduces or eliminates the pain. The downside to this is that the movement of the joint is lost and the patient will never walk normally again [2 4]. The four following principles guide the primary goals for successful ankle fusion: (1) good alignment of the fusion with the hindfoot aligned to the leg and the forefoot to the hindfoot; (2) apposition of broad, flat, vascularized bony surfaces; (3) stable and rigid internal or external fixation; and (4) compression across the arthrodesis site [5]. Numerous techniques for ankle arthrodesis have been described over the years starting as early as 1879 [6]. Currently, there are both open and arthroscopic approaches. The open approaches can be divided into anterior, posterior, medial, lateral, and a combination of medial and lateral (Fig. 47.1). Compared to an arthroscopic approach, the open approach allows for better visualization and access of the joint to correct malalignment; however, there is a greater risk of complications associated with wound healing and soft tissue management than with arthroscopy [7 9]. Further, each approach has a unique set of advantages and disadvantages for visualization, preparing the joint surface, and for strategic placement of internal fixation [10]. The anterior approach allows the best access to the anterior ankle joint;

FIGURE 47.1 Open approaches for ankle arthrodesis: (A) anterior, (B) posterior, (C) lateral, (D) medial, and (E) a combination of medial and lateral. Reproduced with permission from parts (A) and (C) [10]; (B) [13]; (D) and (E) [14].

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FIGURE 47.2 (A) Artist’s depiction and surface anatomy markings to assist in anterior ankle arthroscopy portal placement. (B) Posterior portals and surface anatomy markings showing Achilles tendon and level of talus [15,16]. AL port, Anterolateral portal; AM port, anteromedial portal; LDCB, lateral dorsal cutaneous branch of the SPN; MDCB, medial dorsal cutaneous branch of the SPN; n.v., neurovascular; SPN, superficial peroneal nerve.

however, it lacks exposure to the posterior ankle joint and malleoli. It should be considered when the talus has been displaced medially or laterally under the tibia in the frontal plane. The lateral approach allows for clear visualization of the lateral ankle joint and should be considered when the foot is translated anteriorly as a sequelae of a pilon fracture or when the lateral malleolus must be removed. The medial approach gives excellent access to the anteromedial and posteromedial aspect of the ankle joint. The posterior approach is rarely used as it gives poor visibility to the ankle joint and may possibly lead to inadvertent risk of disturbing the subtalar joint. The combination of medial and lateral approaches provides the greatest visualization and access to the ankle joint. The arthroscopic approach can be divided into anterior and posterior portals (Fig. 47.2) and is inherently less invasive with shorter operative times and less soft tissue injury [11,12]; however, with more challenging visualization and access to the joint, arthroscopic approaches may be more difficult for surgeons with less experience or patients with greater degrees of malalignment. The fusion of the three bones is achieved through different methods of fixation (Fig. 47.3). In general, after the joint has been exposed, the joint is prepared by denuding it of any remaining articular cartilage. Depending on surgeon preference, training, and individual circumstances, screws or screws-and-plate constructs may be used to fuse the ankle in the correct position. Some patients have a concomitant subtalar fusion if there is significant arthritis observed in that joint. There are several methods for internal fixation including screws, plates, and retrograde intramedullary nails. A combination of plates and independent screws are sometimes used. Intramedullary nails are typically reserved for patients that also need their subtalar joint fused. External fixation is another method which is reserved typically for revisions, complex patients with significant bone defects, poor bone quality, and when internal fixation is not feasible, such as in the presence of active or previous infection [17].

47.2.2 Ankle arthroplasty Joint replacement of the ankle has been developed at a slower pace than the other major lower extremity joints (i.e., the knee and hip). This is likely due to the multiple order of magnitudes in difference of the prevalence of these procedures. For instance, in 2010 there were 310,800 total hip replacements and 693,400 total knee replacements for inpatients over 45 years of age [18,19], while in the same year there were less than 1,000 total ankle replacement (TAR) Medicare beneficiaries [20]. The first generation of TARs was attempted in the 1970s but due to poor clinical outcomes, were abandoned until the late 1980s and early 1990s when more anatomically correct second-generation designs were

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FIGURE 47.3 Ankle arthrodesis with independent cannulated screws with anterior ankle fusion plate in several views: (A) lateral, (B) mortise, and (C) anterior. Ankle arthrodesis with independent cannulated screws in several views: (D) lateral, (E) mortise, and (F) anterior.

implemented [21,22]. More modern ankle implant designs have been released nearly continuously since the early 2000s (Figs. 47.4 and 47.5). The first total ankle arthroplasty designs consisted of a long-stem tibial component with an all polyethylene talar body component [25]. However, of the 25 that were implanted, 12 failed and only 7 could be considered a success. Subsequent designs consisted of a polyethylene concave component and a metallic alloy convex component; the systems were either constrained or unconstrained, and bone cement was always used for fixation [21,22]. Throughout the 1970s, a series of implant designs were developed and implemented, but ultimately abandoned—these included the following: Imperial College of London Hospital [26], Newton [27], St. Georg [28], New Jersey Cylindrical Replacement [29,30], Irvine [31], Conaxial [32], Mayo [33 35], Richard Smith [36], Bath-Wessex [37], and the Thompson-Richard prosthesis [38]. Among the reasons these devices failed included bone cement fixation (requiring a large amount of bone resection) [21,39] and over- or under-constraint designs [30,32,35,36,39]. Other complications included: superficial wound healing problems [26,28,40], talar collapse/subsidence [26,28], loosening of the components [26 28,32,35 37], impingement/ankylosis [28], radiolucency [38], and malalignment [40]. The types of failures of these first-generation implants can roughly be grouped into three categories: (1) Failures due to technique: inadequate fixation, poor soft tissue management, malalignment, lack of soft-tissue balancing. (2) Failures of highly-constrained designs: they had excessive transference of shear, compression, and rotatory forces associated with weight bearing to a relatively small area of prosthesis-bone interface; they also required considerable bone resection which made subsidence more common. And (3) Failures of highly-unconstrained designs: less component loosening but unstable and impinged on malleoli or soft tissues. A more comprehensive discussion of the first-generation ankle arthroplasties is available in previous reviews by Vickerstaff et al. [21], Gougoulias et al. [22], or Mulcahy [39]. From the experiences of the first generation of TARs, important design considerations that were incorporated into the second-generation of implants included congruency (i.e., better wear characteristics) and semi-constrained motion (i.e., ability to absorb off axis loading), as well as the limited use of bone cement (i.e., less bone resection) [21,39]. The Agility ankle prosthesis, first implanted in 1984, was the first TAR approved by the FDA and remained the only approved implant until 2005 [39]. While it required the removal of a substantial amount of bone and the fusion of the

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FIGURE 47.4 Second-generation total ankle replacements, such as the (A) STAR, and third-generation total ankle replacements, including the (B) INBONE II and (C) Salto Talaris. STAR, Scandinavian total ankle replacement. Reproduced with permission from parts (A) and (B) [23]; (C) [24].

syndesmosis, in a subsequent long-term follow-up (average time 9 years), more that 90% of the subjects reported decreased pain and satisfaction with the surgical outcome [41]. The Beuchel-Pappas (BP) TAR was another secondgeneration arthroplasty, but with a three-part (i.e., mobile bearing) design [21]. A long-term follow-up found good or excellent results in 88% of 50 ankles, with a 10-year survivorship of 93.5% [42]. Another study considered two designs (shallow and deeps sulcus) of the BP TAR; the 20-year survivorship of 40 shallow-sulcus TARs was 74.2%, while the 12-year survivorship of 75 deep-sulcus designs was 92% [43]. Of note is that this was the first TAR whose long-term results are similar to those found in total knee and hip replacements [21]. The third second-generation TAR is the Scandinavian total ankle replacement (STAR), which is a mobile bearing device that was first implanted in 1981 (Fig. 47.4A) [21]. One study followed 200 patients with STAR TARs for an average of 46 months and found that the 5-year survivorship was 92.7% [44]. Despite these improvements over the first generation of implants, these second generations still failed due to polyethylene wear [39]. As noted earlier, there have been numerous TAR that have come to market in the last two decades. These include third-generation designs such at the INBONE II (Fig. 47.4B) [23] and the Salto Talaris (Fig. 47.4C) [24] and fourthgeneration designs such as the INFINITY (Fig. 47.5A) [23], CADENCE (Fig. 47.5B) [45], and Vantage (Fig. 47.5C) [23]. As many of these devices are still being evaluated, a full review of the third and fourth-generation implants is beyond the scope of this chapter.

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FIGURE 47.5 Fourth-generation total ankle replacements, including the (A) INFINITY, (B) CADENCE, and (C) Vantage. Reproduced with permission from parts (A) and (C) [23]; (C) [24].

47.3 Biomechanical factors in presurgical assessment and consideration of arthroplasty or arthrodesis As noted earlier, there are many aspects for both clinician and patient to consider when deciding whether to treat ESAA with a joint replacement or a fusion. Prior to conducting a TAR, the biomechanics of the ankle must be considered. As described by Bonasia, et al. [46], there are at least three factors to consider: (1) limb and ankle alignment, (2) bony and ligamentous ankle anatomy, and (3) ankle motion in the three cardinal planes. Because arthrodesis is inherently motion limiting, the factors related to motion are not germane to that procedure. However, there are biomechanical concerns that should be considered presurgery—mostly related to bony morphology and alignment.

47.3.1 Limb alignment with arthroplasty The talocrural joint consists of the articulations between the tibia, fibula, and talus; it is often referred to as a mortise (tibia and fibula) and tenon (talus) joint. The superior talar surface is wider anteriorly, so as the ankle dorsiflexes, it “wedges” between medial and lateral malleoli, which increases joint stability [47]. Historically, the ankle has been described as a single degree of freedom hinge that can be represented as a frustrum of a cone with a medial apex whose

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FIGURE 47.6 (A) Ankle mortise view and (B) anterior view of an arthritic ankle joint with coronal plane deformity.

major axis connects the medial and lateral malleoli [48]. However, recent work by Siegler et al. has shown that the ankle should instead be represented as a skewed truncated conic saddle shape with a lateral apex [49]. Another limb alignment concern related to arthroplasty is the initial coronal plane of the ankle (Fig. 47.6). As noted by Kim et al., poor outcomes with TAR have been seen when the accompanying coronal plane malalignment is not addressed surgically [50]. However, even with ankles that have a varus tilt of greater than 10 degrees, the short term (27 months) clinical outcome was comparable to neutral ankles as long as the corrective procedures for the deformity occurred at the same time.

47.3.2 Bony and ligamentous ankle anatomy with arthroplasty The anatomy of the ankle joint is such that it can be represented in two dimensions as a four bar-linkage to describe open kinetic chain dorsiflexion and plantarflexion [51]. Leardini et al.’s research demonstrated that passive ankle joint motion, which can be represented as a four-bar linkage with a single degree of freedom, is determined by the shape of the bones and tension in the ligaments. They concluded that careful consideration of the ankle anatomy must be taken into account with an ankle arthroplasty. A minimum 5-year follow-up on 75 TARs with the Box Ankle (Finsbury Orthopaedics Limited, Leatheread, UK), a “ligament-compatible” ankle replacement, demonstrated significant improvement in function and range of motion, as well as a 97.3% survival rate [52].

47.3.3 Ankle motion in the three cardinal planes with arthroplasty Although the ankle is sometimes considered to have a single degree of freedom, motion does occur in all three planes. During stance phase, the primary motion of the ankle joint is in the sagittal plane; from neutral at heel strike, to 7 degrees of plantarflexion early in stance, to 10 degrees of dorsiflexion in midstance, to 20 degrees of plantarflexion at toe off, before returning to neutral during swing phase [53]. Using a loaded cadaveric model, Michelson and Helgemo found that transverse plane motion was coupled to sagittal plane motion; maximum dorsiflexion led to 2.5 degrees of external rotation, while maximum plantarflexion led to ,1 degree of internal rotation [54]. They also found that maximum dorsiflexion and plantarflexion were associated with ,1 degree of varus.

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47.3.4 Bony morphology and alignment with arthrodesis When treating a varus foot deformity with arthrodesis, the cause (bone deformity, chronic ligament insufficiency, muscle imbalance, or a combination) and the location (hindfoot, midfoot, or forefoot) of the deformity must be considered [55]. For instance, a physically active, heavy manual laborer might be a candidate for a fusion; this would be further indicated if the deformity is very severe and located in the hindfoot. For any ankle arthrodesis, choice of surgical techniques depends on the underlying cause. While external fixators are preferred for subjects with sepsis or osteopenia, the choice between arthroscopic techniques or an open fusion depends on the amount of deformity, with the former indicated for mild cases and the latter for more severe [56]. Other groups have demonstrated the importance of presurgical alignment. In a retrospective study of 215 uncomplicated open ankle fusions, Chalayon et al. found that preoperative varus ankle alignment was a significant risk factor for nonunion [57].

47.4

Biomechanical considerations/complications of arthroplasty or arthrodesis

There are many potential biomechanical considerations/complications from treating ESAA patients with an ankle arthrodesis or arthroplasty, including some which are patient related, such as ankle alignment, gait mechanics, and arthritis at distal joints, and others that are laboratory-based, like component wear or failure simulators, cadaveric gait simulators, and computational modeling.

47.4.1 Ankle alignment/malalignment Early work by Hefti et al. demonstrated that ankle arthrodesis should be performed in a neutral sagittal plane position [58]. Mann suggested “An ankle arthrodesis must be carefully aligned into slight valgus and neutral dorsiflexion, plantarflexion and the same degree of external rotation as the opposite leg [59].” Similarly, Buck et al. found that to obtain normal knee function, the ankle should be fused at neutral flexion, 0 5 degrees valgus, and 5 10 degrees external rotation [60]. A study by Hintermann et al. noted that of 30 ankles that were converted from a painful arthrodesis to an arthroplasty, 28 had at least a 10-degree malalignment in one or more planes [61]. Mild varus/valgus malalignment has been shown to alter the ankle joint contact mechanics [62].

47.4.2 Gait mechanics Recently, Deleu et al. have published a metaanalysis on the change in gait mechanics after ankle arthrodesis or ankle arthroplasty [63]. They found 17 studies encompassing 883 ESAA patients with pre- and postsurgical kinematics, kinetics, and spatio-temporal measurements. There was moderate evidence of improvements in gait biomechanics, including ankle kinetic variables such as peak plantarflexion moment and ankle power generation after arthroplasty, but less so with arthrodesis, which was limited to increased ankle dorsiflexion moments. The main benefits of arthroplasty over arthrodesis include conservation of presurgical ankle joint range of motion and the protection of other lower limb joints. Both surgical groups walked faster postoperatively. However, as noted by the authors “The currently available evidence base of research papers evaluating changes in gait biomechanics after TAR and ankle arthrodesis is limited by a lack of prospective research, low sample sizes and heterogeneity in the patho-etiology of ankle osteoarthritis.”

47.4.3 Arthritis at distal joints Ankle arthroplasty and arthrodesis both fundamentally affect the biomechanics of the foot and ankle, but quantifying the effects can be difficult. In a retrospective study of 1001 primary TARs, Gross et al. found that 26 patients (2.6%) required either a subtalar (18), talonavicular (3), subtalar and talonavicular (3), or a triple (subtalar, talonavicular, and calcaneocuboid) arthrodesis [64]. This study does not demonstrate causation, but suggests that TAR can affect the distal foot joints (Fig. 47.7). A radiographic analysis of 197 TAR patients with both fixed and mobile-bearing implants found that the sagittal plane range of motion through the prosthesis was 68% of the total range of motion from the tibia to the ground [65]. In other words, 32% of the motion occurred at the peritalar joints, perhaps indicating increased midfoot and subtalar motion. This same group also examined 140 TAR patients with an average follow-up of 6.5 years and found radiographic evidence of subtalar (27%) and talonavicular (31%) arthritis, suggesting a moderate but nominal increase in arthritis at adjacent joints [66]. In a mid-term follow-up study of 941 primary TAR patients it was found that in 4% (37) a secondary subtalar joint fusion was necessary [67]. On a subset of 671 patients, a secondary

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FIGURE 47.7 Subtalar arthrosis after an ankle joint fusion.

radiographic analysis found that 39% (260) of patients had minor subtalar osteoarthritic changes while 60% had major osteoarthritic changes. Concerning arthrodesis, one retrospective study of 82 ankles at a mean follow-up of 47 months demonstrated that subsequent fusions were required in 7 feet (5 subtalar, 1 talonavicular/calcaneocuboid, and 1 triple) [68]. Another study conducted a radiographic analysis on 23 patients who had an isolated ankle arthrodesis with a mean follow-up of 22 years [69]. OA of all ipsilateral foot joints (subtalar, talonavicular, calcaneocuboid, naviculocuneiform, tarsometatarsal, and first metatarsophalangeal) was more severe than the OA of those joints on the contralateral side. Another long-term study retrospectively examined 18 arthrodeses with a mean follow-up of 23 years [70]. Using a 0 to 4 scale of degenerative arthritis, they found a mean grade of 3.1 6 0.54 for the subtalar joint and 2.3 6 0.83 for the talonavicular joint, indicating that arthritis was developing at the ipsilateral joints. Finally, a medium- to long-term retrospective study of 66 eligible ankle arthrodeses (mean follow-up of 9 6 4.1 years) demonstrated significant arthritis at all contiguous (subtalar, talonavicular and calcaneocuboid) joints [71].

47.4.4 Component wear or failure As ankle arthroplasties become more wide spread with increased longevity, wear effects must be considered, in particular with TARs that use mobile bearings. Often times, wear simulators are employed to study these phenomena. For example, Smyth et al. found that wear rates were highly dependent on the inclusion of internal/external rotation in the wear simulation profile [72]. However, a recent review by Mujukian et al. found a wide range of magnitudes in the kinematic and kinetic parameters, with only two studies conducting any validation with independently derived data [73]. They concluded that wear simulators may not accurately represent the in vivo wear of TARs. Rather than using a wear simulator, Espinosa et al. investigated wear due to implant misalignment using two validated computer models [74]. They found that highly congruent designs yield lower contact stresses and that malalignments of greater than 5 degrees can cause stresses above the yield stress of polyethylene.

47.4.5 Cadaveric gait simulation of arthroplasty and arthrodesis Conducting cadaveric gait simulations of ankle fusions or total joint replacements has distinct limitations, but also allows for experiments that could not be conducted on living subjects. The primary limitation is that healing postsurgery can obviously not be simulated. Therefore any cadaveric-based simulation of arthroplasty or arthrodesis is only able to explore immediate postoperative strength. Furthermore, loading across these constructs is limited, as much like living humans, full body weight cannot safely be borne immediately postsurgery. That said, useful knowledge is still obtainable. Of particular interest is the loading across the ankle joint post arthroplasty. A group at the Hospital for Special Surgery conducted a two-part study, whereby cadaveric gait simulation was first employed with a musculoskeletal (OpenSim) foot and ankle model to determine the peak compressive loads and moments across the ankle [75]; the second part, which is summarized in the next section, explored the loading on the implant itself. Simulations were conducted at one-quarter body weight and one-sixth speed. This combined musculoskeletal model and cadaveric gait simulation was contrasted to previous work that involved two or three-dimensional static models, or skin-mounted

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marker-based modeling, or finite element simulations of a few loading conditions, demonstrating the unique opportunities to evaluate the pre- and postoperative conditions. Cadaveric gait simulation has also proven very useful in studying the effect of malalignment in both arthrodesis and arthroplasty (Fig. 47.8). In a six-part study, the group at the VA Puget Sound has explored the effect of anterior/posterior talar translation and malalignment in the frontal (inversion/eversion) and transverse (internal/external rotation) planes for both ankle fusions and total joint replacements using a robotic gait simulator. Only 3 of the 6 studies have been published to date. Anterior/posterior translation with ankle arthrodesis affected the joint kinematics of several distal joints, most significantly when the displacement was 6 mm or greater in the posterior direction [76]. Similarly, when investigating anterior/posterior translation with ankle arthroplasty, it was found that posterior malalignment resulted in more differences than anterior malalignment, with the calcaneal cuboid joint being the most affected [77]. Finally, frontal plane malalignment of ankle arthroplasty was also studied. Due to a small number of specimens, the findings were not strong. However, the results indicated a trend that varus malalignment leads to valgus and external rotation of the subtalar and talonavicular joints [78].

47.4.6 Computational models of arthroplasty and arthrodesis As computer power becomes more efficient and biomechanical models become more sophisticated, computational models of ankle OA treatments are becoming more common. Finite element (FE) models have been developed to study polyethylene wear rates in response to changes in pressure due to joint malignment [74], indicating the malalignments as small as 5 degrees that lead to loading above the yield stress of polyethylene. Other, more detailed FE foot models have compared and contrasted both ankle fusion and ankle joint replacement with normal feet [79]. Output parameters included plantar pressure, joint contact pressure, and bone stress. This model demonstrated that, for example, there was not an increase in subtalar joint stress for either ESAA treatment. More recently, FE foot models have been used to investigate the interface between the total joint replacement and the bone [80]. Implant micromotion was studied for various implant designs and loading conditions, and the importance of considering patient specific anatomy and complex loading patterns was demonstrated.

FIGURE 47.8 (A) Representative preoperative lateral radiograph of a cadaveric TAR with loading jig (a) and scale bar (b). The TTR, a measure of anterior/posterior misalignment, was calculated from four cortical surfaces identified in each radiograph. Yellow arcs represent the four cortical surfaces fit with circles (left to right): the anterior aspect of the talus, the superimposed talar domes, the posterior subtalar articular surface, and the posterosuperior cortex of the calcaneus. The talar axis was drawn parallel to the plantar surface of the foot through the intersection of the subtalar and posterosuperior calcaneal contours to the most anterior point of the talus, identified as a vertical projection of the anterior talar contour. The tibial axis was drawn through the midpoint of the tibial shaft at 5 and 10 cm above the talar dome. The intersection of the tibial and talar axes divides the talar axis into two lengths, X and Y. The TTR was calculated as X/(X 1 Y). An anteriorly displaced talus results in a larger TTR value. (B) Representative postoperative radiograph with the talar component at the 9 A malalignment and (C) the 9 P malalignment. TAR, Total ankle replacement; TTR, tibiotalar ratio [73].

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Biomechanical outcomes

There are dozens of studies in the literature that evaluate a series of patients undergoing arthroplasty or arthrodesis with respect to biomechanical outcomes. Few exist that compare the two treatments. There are several biomechanical considerations after surgery for ESAA such as stride length and cadence, range of motion in the sagittal and coronal planes, strength and power particularly in the sagittal plain, heel strike forces, and actual step activity. These may or may not be associated with an improvement in patient-reported function or decrease in pain; however, all are important in terms of restoring the patient’s ankle to its original biomechanical function which may impact other important domains such as implant wear and future adjacent joint arthritis. Among the studies that have compared biomechanical outcomes between the two treatments, ankle arthroplasty has demonstrated superior outcomes in terms of overall arc of motion [2,81], sagittal plain range of motion [3,4,81,82], sagittal power [3], step counts [83], and restoration of gait to as close to normal as possible. Those studies that have evaluated overall gait function have concluded that neither restore the patient back to normal gait, but gait after ankle arthroplasty most closely resembles a normalized pattern [4,81]. While sagittal motion is less after arthrodesis, studies have found an increase in hip motion, coronal motion, and talonavicular motion compared to arthroplasty which may predispose fusion patients to more degenerative changes in other joints over time [3,81]. Peak plantarflexion and heel strike forces have also been greater after arthrodesis [2,82]. Studies comparing step activity using step activity monitors are rare [2]; however, one found that improvement after arthrodesis and ankle replacement may follow different trajectories. TAR patients showed more improvement than arthrodesis patients early in recovery. TAR patients demonstrated rapid improvement over the first 6 12 months postop while arthrodesis patients had little to no improvement. However, after 3 years postop both groups showed similar improvement. This study also found that step activity was associated with several patient-reported outcomes but the improvements were not in parallel. Patient-centered outcomes tend to improve in the acute phase and often plateau in the first 6 12 months while step activity continues to improve. Little has been studied on comparing the postoperative gait of these two procedures, especially on uneven surfaces. One study reported on 77 consecutive patients (61 arthroplasty and 16 arthrodesis) who completed 12 months of follow-up [84]. Both groups improved in their performance walking on uneven surfaces; however, arthroplasty patients had a significantly better outcome than the arthrodesis patients in walking upstairs, downstairs, and uphill.

47.6

Clinical outcomes

Until recently, most surgeons have considered ankle arthrodesis to be the gold standard, and studies have shown that 80% to 85% of individuals have successful outcomes [85]. Even in long-term studies, an average of 9 years postarthrodesis, fusion has been achieved in as high as 91% of patients, and a comparable percentage reported satisfaction with their procedure [71]. While ankle arthrodesis has a long track record of being an effective treatment for ESAA, fusing joints comes with a cost. Challenges include the extended time required to achieve fusion, potential for nonunion, leg length discrepancy, malalignment, stress fractures, and continued pain [9,85]. One of the most significant drawbacks is that many patients eventually develop arthritis in distal joints (e.g., the subtalar and midtarsal) [69,86,87], which can require secondary procedures [88]. This leads to the strong consideration for ankle arthroplasty where ankle motion is retained. When considering clinical outcomes for any operative procedure, the four most important outcome domains include safety, effectiveness, costs, and patient subgroups. That is, are there subsets of patients that respond more favorably to one treatment over another? With respect to ankle arthrodesis and arthroplasty, evidence is emerging that may help clarify some of these questions, but many unknowns and uncertainties still exist. For example, patients who are younger or heavier tend to undergo arthrodesis while those who are older and lighter tend to undergo arthroplasty despite little to no evidence in the literature to support these decisions. Surgeons are left to make decisions based on their personal experience or clinical gestalt. This section will briefly summarize what we know with respect to these four outcome domains.

47.6.1 Safety Survival of implants (Fig. 47.9) and failure of fusion (Fig. 47.10) are often considered the most important concerns for these two treatments. Although results were disappointing for the first-generation of ankle prostheses [89], a second generation was developed in the 1990s that has shown more promising results. These were better able to reproduce normal ankle anatomy, stability, alignment, and joint kinematics, with a mean survivorship rate of 91% at 5-year, 83% at 10-year and 66% at 15-year follow-up [90]. Third-generation prostheses, which were introduced in the 2000s, were designed to

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FIGURE 47.9 Salto Tolaris total ankle arthroplasty that has failed due to osteolysis; (A) anterior/posterior view and (B) mortise view.

FIGURE 47.10 Ankle arthrodesis that has failed due to a nonunion; (A) anterior/posterior view and (B) mortise view.

facilitate even greater ankle mobility [42,91,92], with a mean survivorship rate of 93% at 5-years and 83% at 10-year follow-up [90]. More recent studies report high revision rates after arthroplasty. A recent study of 74 patients with posttraumatic arthrosis were treated using arthroplasty with a Tornier Salto prosthesis, a third-generation prosthesis [93]. Sixty of these patients achieved a mean 59-month follow-up and experienced a 12-month, 24-month, and overall revision rate of 8%, 18%, and 42%, respectively. The most commonly performed procedures were cyst debridement and autologous spongy bone grafting (20%). Fifteen percent (9) of the prosthetics were explanted or switched to a tibiotalar arthrodesis. Patients who underwent revision had worse outcomes compared to those who did not have a revision. A recent short-term follow-up study suggested ankle-specific adverse events were infrequent with little difference between arthroplasty and arthrodesis; however, ankle-specific adverse events negatively impacted patient improvement compared to those with no AEs regardless of treatment approach [94]. So clearly adverse events or complications impact future function and should be minimized. A previous study found that survival rates after arthroplasty were significantly higher after the first 30 procedures that the surgeon performed [95], suggesting experience is a factor which should be considered.

47.6.2 Effectiveness Several systematic reviews [8,9,96 98] suggest that short-term clinical and functional outcomes appear similar if not superior after arthroplasty, but reoperation rates or major complications are greater. A recent metaanalysis [96],

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evaluating 10 studies, concluded similar outcomes especially in short-term follow-up and confirmed an increase in arthroplasty complications. These systematic reviews report many study limitations and recommend further high quality research comparing these two treatments over a longer follow-up period. A study evaluating 273 consecutive patients (103 arthrodesis and 170 arthroplasty) with 3-year patient-centered outcomes demonstrated results favoring arthroplasty [99]. Patients reported improved pain reduction and function after both surgical treatments with improvements in the Musculoskeletal Functional Assessment and SF-36 Physical Function scores favoring arthroplasty, particularly when the arthroplasty was done with later generation implants. The improvements were also greater in younger patients. A larger multisite prospective cohort study including 517 participants (414 arthroplasty and 103 arthrodesis) with 4-year patient outcomes reported significant improvement after both procedures in the Foot and Ankle Ability Measure sports and activities of daily living scores, SF-36 physical component summary scores, and all pain measures. Improvements were significantly greater after arthroplasty in nearly every patient-reported measure. These differences were clinically significant, particularly in the ankle-specific measures. Further, 28% more arthroplasty patients were completely satisfied than those who underwent arthrodesis. Revision rates were 8.7% and 17.5% in the arthroplasty and arthrodesis groups, respectively [100].

47.6.3 Costs One cost benefit analysis comparing arthroplasty, arthrodesis, and nonoperative management used Markov model analysis from a health-systems perspective [101]. Costs were derived from the 2012 Nationwide Inpatient Sample and expressed in 2013 US dollars. Effectiveness was expressed in quality-adjusted life years (QALYs). The primary outcome measure was the incremental cost-effectiveness ratio. When indirect costs were included, arthroplasty was both more effective and resulted in $5,900 and $800 in lifetime cost savings compared with the lifetime costs following nonoperative management and arthrodesis, respectively. At a $100,000/QALY threshold, surgical management of ankle arthritis was preferred for patients younger than 96 years, and arthroplasty was increasingly more cost-effective in younger patients. Arthroplasty, arthrodesis, and nonoperative management were the preferred strategy in 83%, 12%, and 5% of the analyses, respectively. The authors concluded that as indications for and utilization of arthroplasty increase, continued research is needed to define appropriate subgroups of patients who would likely derive the greatest clinical benefit from that procedure.

47.6.4 Patient subgroups In the treatment of ESAA after failed conservative care, many studies are limited by small sample sizes, varying inclusion criteria, lack of control for differences between treatment groups or surgeon experience, short follow-up, and weak patient-centered outcome measures. In addition, only one study exists comparing ankle arthroplasty to ankle arthrodesis that evaluated whether certain patient subgroups respond more favorably to one treatment or another. Studies evaluating the comparative effectiveness of these treatments, are based on average improvements and do not account for heterogeneity of treatment effects. Until there is a better understanding of which patients respond more favorably to one treatment versus another, surgeons have to base their recommendations on averages and their clinical experience. One study has been published that did attempt to evaluate the differential effects of these two treatments in the same comparative study [102]. Treatment success was defined as a patient who achieved their minimal clinically important difference of nine points in the Foot and Ankle Ability Measure Activities of Daily Living score and did not experience a minor or major revision. The 24-month treatment success rate was significantly higher (p 5 0.016) for the arthroplasty group (81%, 95% CI 76% 84%) than for the arthrodesis group (68%, 95% CI 58% 77%), OR 1.9 (95% CI 1.1, 3.2), adjusting for age, sex, BMI, and the Functional Comorbidity Index (FCI). However, this difference was only observed in specific subgroups. In those patients with higher FCI scores (FCI 5 4), 80% of arthroplasty patients experienced treatment success versus 62% of arthrodesis patients. In those with lower scores (FCI 5 2), treatment success rates were 81% and 77%, respectively. In other words, patients with a higher comorbidity profile fared better after arthroplasty but there was little difference in those with lower comorbidity profiles. Similarly, 81% of arthroplasty patients who were not employed full-time experienced treatment success versus just 58% who underwent arthrodesis. Success rates were similar in those who were fully employed with treatment success rates of 79% and 78%, respectively. In other words, part time workers fared better after arthroplasty but there was no difference in those who worked full time. There were no significant differential treatment effects for age, BMI, or sex, suggesting that the traditional assumption that patients who are heavier and younger should undergo fusion rather than arthroplasty is incorrect. However, with only a 2-year follow-up, these subgroup effects should be considered with caution and more research on the heterogeneity of treatment effects for arthroplasty and arthrodesis should be evaluated with a long-term follow-up.

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Areas of future biomechanical research

Understanding and quantifying the functional differences between ankle fusions and total joint replacements remain a fundamental area of research. As more and more third- and fourth-generation TARs are developed and implanted, the effect of these devices on the joints distal to the ankle remains unknown. It is assumed that a TAR will result in more normal joint kinematics for the foot joints distal to the ankle as compared to an ankle fusion. However, this has not been quantified and detailed foot bone kinematics studies conducted with biplane fluoroscopy are required. Furthermore, longer term follow-up of subtalar and talonavicular joint function after TAR or arthrodesis are necessary to understand the differences in patient outcomes.

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The Clinical Biomechanics Award 2013—presented by the International Society of Biomechanics: new observations on the morphology of the talar dome and its relationship to ankle kinematics. Clin Biomech (Bristol, Avon) 2014;29(1):1 6. [50] Kim BS, Choi WJ, Kim YS, Lee JW. Total ankle replacement in moderate to severe varus deformity of the ankle. J Bone Jt Surg Br 2009;91 (9):1183 90. [51] Leardini A, O’Connor JJ, Catani F, Giannini S. A geometric model of the human ankle joint. J Biomech 1999;32(6):585 91. [52] Giannini S, Romagnoli M, Barbadoro P, Marcheggiani Muccioli GM, Cadossi M, Grassi A, et al. Results at a minimum follow-up of 5 years of a ligaments-compatible total ankle replacement design. Foot Ankle Surg 2017;23(2):116 21. [53] Perry J. Gait analysis: normal and pathological function. Thorofare, NJ: SLACK Incorporated; 1992. [54] Michelson JD, Helgemo Jr. SL. Kinematics of the axially loaded ankle. Foot Ankle Int 1995;16(9):577 82. [55] AlSayel F, Valderrabano V. Arthrodesis of a varus ankle. Foot Ankle Clin 2019;24(2):265 80. [56] Abidi NA, Gruen GS, Conti SF. Ankle arthrodesis: indications and techniques. J Am Acad Orthop Surg 2000;8(3):200 9. [57] Chalayon O, Wang B, Blankenhorn B, Jackson 3rd JB, Beals T, Nickisch F, et al. Factors affecting the outcomes of uncomplicated primary open ankle arthrodesis. Foot Ankle Int 2015;36(10):1170 9. [58] Hefti FL, Baumann JU, Morscher EW. Ankle joint fusion—determination of optimal position by gait analysis. Arch Orthop Trauma Surg 1980;96(3):187 95. [59] Mann RA. Surgical implications of biomechanics of the foot and ankle. Clin Orthop Relat Res 1980;146:111 18. [60] Buck P, Morrey BF, Chao EY. The optimum position of arthrodesis of the ankle. A gait study of the knee and ankle. J Bone Jt Surg Am 1987;69(7):1052 62. [61] Hintermann B, Barg A, Knupp M, Valderrabano V. Conversion of painful ankle arthrodesis to total ankle arthroplasty. Surgical technique. J Bone Jt Surg Am 2010;Suppl. 1(Pt 1):55 66 92. [62] Wayne JS, Lawhorn KW, Davis KE, Prakash K, Adelaar RS. The effect of tibiotalar fixation on foot biomechanics. Foot Ankle Int 1997;18 (12):792 7. [63] Deleu PA, Besse JL, Naaim A, et al. Change in gait biomechanics after total ankle replacement and ankle arthrodesis: a systematic review and meta-analysis. Clin Biomech (Bristol, Avon) 2020;73:213 25. [64] Gross CE, Lewis JS, Adams SB, Easley M, DeOrio JK, Nunley 2nd JA. Secondary arthrodesis after total ankle arthroplasty. Foot Ankle Int 2016;37(7):709 14. [65] Dekker TJ, Hamid KS, Easley ME, DeOrio JK, Nunley JA, Adams Jr. SB. Ratio of range of motion of the ankle and surrounding joints after total ankle replacement: a radiographic cohort study. J Bone Jt Surg Am 2017;99(7):576 82.

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[66] Dekker TJ, Walton D, Vinson EN, Hamid KS, Federer AE, Easley ME, et al. Hindfoot arthritis progression and arthrodesis risk after total ankle replacement. Foot Ankle Int 2017;38(11):1183 7. [67] Sokolowski M, Krahenbuhl N, Wang C, et al. Secondary subtalar joint osteoarthritis following total ankle replacement. Foot Ankle Int 2019;40 (10):1122 8. [68] Gordon D, Zicker R, Cullen N, Singh D. Open ankle arthrodeses via an anterior approach. Foot Ankle Int 2013;34(3):386 91. [69] Coester LM, Saltzman CL, Leupold J, Pontarelli W. Long-term results following ankle arthrodesis for post-traumatic arthritis. J Bone Jt Surg Am 2001;83-A(2):219 28. [70] Fuchs S, Sandmann C, Skwara A, Chylarecki C. Quality of life 20 years after arthrodesis of the ankle. A study of adjacent joints. J Bone Jt Surg Br 2003;85(7):994 8. [71] Hendrickx RP, Stufkens SA, de Bruijn EE, Sierevelt IN, van Dijk CN, Kerkhoffs GM. Medium- to long-term outcome of ankle arthrodesis. Foot Ankle Int 2011;32(10):940 7. [72] Smyth A, Fisher J, Suner S, Brockett C. Influence of kinematics on the wear of a total ankle replacement. J Biomech 2017;53:105 10. [73] Mujukian A, Ho NC, Day MJ, Ebramzadeh E, Sangiorgio SN. A systematic review of unsystematic total ankle replacement wear evaluations. JBJS Rev 2020;8(3):e0091. [74] Espinosa N, Walti M, Favre P, Snedeker JG. Misalignment of total ankle components can induce high joint contact pressures. J Bone Jt Surg Am 2010;92(5):1179 87. [75] Steineman BD, Quevedo Gonzalez FJ, Sturnick DR, Deland JT, Demetracopoulos CA, Wright TM. Biomechanical evaluation of total ankle arthroplasty. Part I: joint loads during simulated level walking. J Orthop Res 2021;39(1):94 102. [76] Imsdahl SI, Stender CJ, Cook BK, Pangrazzi G, Patthanacharoenphon C, Sangeorzan BJ, et al. Anteroposterior translational malalignment of ankle arthrodesis alters foot biomechanics in cadaveric gait simulation. J Orthop Res 2019. [77] McKearney DA, Stender CJ, Cook BK, Moore ES, Gunnell LM, Monier BC, et al. Altered range of motion and plantar pressure in anterior and posterior malaligned total ankle arthroplasty: a cadaveric gait study. J Bone Jt Surg Am 2019;101(18):e93. [78] Buckner BC, Stender CJ, Baron MD, Hornbuckle JHT, Ledoux WR, Sangeorzan BJ. Does coronal plane malalignment of the tibial insert in total ankle arthroplasty alter distal foot bone mechanics? A cadaveric gait study. Clin Orthop Relat Res 2020;478(7):1683 95. [79] Wang Y, Wong DW, Tan Q, Li Z, Zhang M. Total ankle arthroplasty and ankle arthrodesis affect the biomechanics of the inner foot differently. Sci Rep 2019;9(1):13334. [80] Quevedo Gonzalez FJ, Steineman BD, Sturnick DR, Deland JT, Demetracopoulos CA, Wright TM. Biomechanical evaluation of total ankle arthroplasty. Part II: influence of loading and fixation design on tibial bone-implant interaction. J Orthop Res 2021;39(1):103 11. [81] Pedowitz DI, Kane JM, Smith GM, Saffel HL, Comer C, Raikin SM. Total ankle arthroplasty vs ankle arthrodesis: a comparative analysis of arc of movement and functional outcomes. Bone Jt J 2016;98-B(5):634 40. [82] Segal AD, Cyr KM, Stender CJ, Whittaker EC, Hahn ME, Orendurff MS, et al. A three-year prospective comparative gait study between patients with ankle arthrodesis and arthroplasty. Clin Biomech (Bristol, Avon) 2018;54:42 53. [83] Shofer JB, Ledoux WR, Orendurff MS, Hansen ST, Davitt J, Anderson JG, et al. Step activity after surgical treatment of ankle arthritis. J Bone Jt Surg Am 2019;101(13):1177 84. [84] Jastifer J, Coughlin MJ, Hirose C. Performance of total ankle arthroplasty and ankle arthrodesis on uneven surfaces, stairs, and inclines a prospective study. Foot Ankle Int 2014. p. 1071100714549190. [85] Muir DC, Amendola A, Saltzman CL. Long-term outcome of ankle arthrodesis. Foot Ankle Clin 2002;7(4):703 8. [86] Thomas R, Daniels TR, Parker K. Gait analysis and functional outcomes following ankle arthrodesis for isolated ankle arthritis. J Bone Jt Surg Am 2006;88(3):526 35. [87] Beyaert C, Sirveaux F, Paysant J, Mole D, Andre JM. The effect of tibio-talar arthrodesis on foot kinematics and ground reaction force progression during walking. Gait Posture 2004;20(1):84 91. [88] Buechel FF, Pappas MJ, Iorio LJ. New Jersey low contact stress total ankle replacement: biomechanical rationale and review of 23 cementless cases. Foot Ankle Int 1988;8(6):279 90. [89] Martin RL, Stewart GW, Conti SF. Posttraumatic ankle arthritis: an update on conservative and surgical management. J Orthop Sports Phys Ther 2007;37(5):253 9. [90] Roukis TS, Bartel AF. Survivorship of first-, second-, and third-generation total ankle replacement systems. Primary and revision total ankle replacement. Springer; 2016. p. 15 23. [91] Kofoed H, Lundberg-Jensen A. Ankle arthroplasty in patients younger and older than 50 years: a prospective series with long-term follow-up. Foot Ankle Int 1999;20(8):501 6. [92] Pyevich MT, Saltzman CL, Callaghan JJ, Alvine FG. Total ankle arthroplasty: a unique design. two to twelve-year follow-up. J Bone Jt Surg Am 1998;80(10):1410 20. [93] Gramlich Y, Neun O, Klug A, Buckup J, Stein T, Neumann A, et al. Total ankle replacement leads to high revision rates in post-traumatic endstage arthrosis. Int Orthop 2018;42(10):2375 81. [94] Norvell DC, Shofer JB, Hansen ST, Davitt J, Anderson JG, Bohay D, et al. Frequency and impact of adverse events in patients undergoing surgery for end-stage ankle arthritis. Foot Ankle Int 2018. p. 1071100718776021. ˚ . The Swedish ankle arthroplasty register: an analysis of 531 arthroplasties between 1993 and 2005. Acta [95] Henricson A, Skoog A, Carlsson A Orthop 2007;78(5):569 74.

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[96] Kim HJ, Suh DH, Yang JH, Lee JW, Kim HJ, Ahn HS, et al. Total ankle arthroplasty vs ankle arthrodesis for the treatment of end-stage ankle arthritis: a meta-analysis of comparative studies. Int Orthop 2017;41(1):101 9. [97] Lawton CD, Butler BA, Dekker 2nd RG, Prescott A, Kadakia AR. Total ankle arthroplasty vs ankle arthrodesis-a comparison of outcomes over the last decade. J Orthop Surg Res 2017;12(1):76. Available from: https://doi.org/10.1186/s13018-017-0576-1. [98] Maffulli N, Longo UG, Locher J, Romeo G, Salvatore G, Denaro V. Outcome of ankle arthrodesis and ankle prosthesis: a review of the current status. Br Med Bull 2017;124(1):91 112. [99] Benich MR, Ledoux WR, Orendurff MS, Shofer JB, Hansen ST, Davitt J, et al. Comparison of treatment outcomes of arthrodesis and two generations of ankle replacement implants. J Bone Jt Surg Am 2017;99(21):1792 800. [100] Sangeorzan BJ, Ledoux WR, Shofer JB, Davitt J, Anderson JG, Bohay D, et al. Comparing 4-year changes in patient-reported outcomes following ankle arthroplasty and arthrodesis. J Bone Jt Surg Am 2021;103(10):869 78. [101] Nwachukwu BU, McLawhorn AS, Simon MS, Hamid KS, Demetracopoulos CA, Deland JT, et al. Management of end-stage ankle arthritis: cost-utility analysis using direct and indirect costs. J Bone Jt Surg Am 2015;97(14):1159 72. [102] Norvell DC, Ledoux WR, Shofer JB, Hansen ST, Davitt J, Anderson JG, et al. Effectiveness and safety of ankle arthrodesis vs arthroplasty: a prospective multicenter study. J Bone Jt Surg Am 2019;101(16):1485 94.

Chapter 48

Prosthetic Feet Glenn K. Klute1,2 1

Center for Limb Loss and MoBility (CLiMB), Department of Veterans Affairs, Seattle, WA, United States, 2Department of Mechanical Engineering,

University of Washington, Seattle, WA, United States

Abstract Prosthetic feet are prescribed to individuals with major lower limb amputations who are capable of ambulation. Of the thousands of steps taken each day on prosthetic feet, one in five are other than in a straight line on level ground. Different prosthetic features are used to accommodate the spectrum of ambulatory activities. Solid ankle cushion heel prosthetic feet are intended for limited household ambulation. Fixed-angle stiffness feet enable much broader function and are widely prescribed. Variable-angle stiffness feet accommodate sloping terrain; while yet to be commercialized, variable stiffness feet are adaptable to a wide variety of activities. Powered prosthetic feet offer promise, but further development is needed to optimize performance and prescription practice. This chapter describes the demographics of individuals with lower limb amputation, what is known about how prosthetic feet are used, the form and function of commercially available prosthetic feet, and prognostication on future directions.

48.1

Introduction

A prosthetic foot is an essential component of every prosthesis prescription for ambulatory individuals with a major lower limb amputation. Data compiled over the last two decades can provide some insight into those who might receive a prosthesis prescription. In the United States, the population of 332 million [1] has a median age of 37 years, is 51% female [2] and 11% diabetic [3]. Approximately 623,000 of this population live with a major lower limb amputation [4], the primary cause of which is related to compromised vasculature, particularly due to diabetes [5]. This is a reason for concern in an aging and overweight population as pooled mortality rates of those who underwent nontraumatic major lower limb amputation were 34% at their 1-year and 64% at their 5-year follow-ups, with diabetic patients exhibiting 27% at their 1-year and 63% at their 5-year follow-ups [6]. Amputation level is a significant factor in survival, with mortality rates for 1- and 5-year follow-ups for transfemoral amputation at 50% and 78%, respectively, much worse than the 1- and 5-year follow-ups for transtibial amputation at 26% and 62%, respectively [7]. While females comprise half the US population, only about one-third of those with limb loss are female [4,8]. Within the general US population, two interesting cohorts arise among those with lower limb amputation. The US veteran population of 18 million has a median age of 65 years [9] of whom only 8% are female [10], but 21% are diabetic [11]. About 9 million of these veterans receive care from the VA at one of its medical centers or outpatient sites [12]. Available data indicates the VA performs over 6000 amputations per year, of which 37% are major limbs and 1.8% involve females [13]. The veteran population is older with additional comorbidities, suggesting the mortality risk is perhaps worse than the US general population. The other cohort of individuals with lower limb amputation in the US stems from military conflicts overseas. This population is significantly smaller (i.e., n 5 1,795 major limb amputation (all causes) from 2000 to 2011 [14]; n 5 1,645 major limb amputation (battle injuries) from 2001 to 2015 [15]) and were younger at the time of amputation, with 51% between 20 and 24 years of age, 25% between 25 and 29 years of age, and 19% 30 years of age and older [14]. Importantly, these individuals met military fitness standards at the time of their amputation. Comparing these two cohorts, the veteran population is older with additional comorbidities with undesirable mortality risk, whereas the cohort arising from recent military conflicts may be expected to need rehabilitative care for many years. For individuals who have the potential to ambulate following lower limb amputation, a key rehabilitation milestone is the successful fitting of a prosthesis [16]. The prosthesis is used to achieve functional mobility, and greater prosthesis Foot and Ankle Biomechanics. DOI: https://doi.org/10.1016/B978-0-12-815449-6.00027-5 © 2023 Elsevier Inc. All rights reserved.

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use can lead to higher levels of function, quality of life, and independence due to improved mobility [17]. However, not all individuals with a lower limb amputation are candidates for prosthesis prescription. Only about one-third of US veteran amputees receive a prosthetic prescription [18], but prescription rates vary by amputation level. Forty-three percent of veterans with a transtibial amputation were fit with a prosthesis within one year, while only 20% with a transfemoral amputation were fit.

48.2

Prescription and expected use of prosthetic feet

For individuals with lower limb loss and the potential to ambulate, the clinician must choose among several hundred available prosthetic feet while trying to optimize function, independence, and quality of life for the individual. At one end of the function spectrum is the need to provide safety and body weight support to the household ambulator during transfers and other, minimalist activities of daily living. At the other end is the elite athlete for whom maximizing speed is an important consideration. For most individuals in the US, prosthesis prescription is based on the Medicare Local Coverage Determination for Lower Limb Prostheses [19]. While half have Medicare or Medicaid as the primary payer [8], many private insurers also use the Medicare functional classification levels (MFCL) for billing purposes. Individual eligibility for a covered lower limb prosthesis is indicated when the person will reach or maintain a defined functional level within a reasonable time and is motivated to ambulate. The clinical determination of potential MFCL is based on the reasonable expectations of the prosthetist and treating practitioner and includes the individual’s history, their current condition, and their desire to ambulate. The five MFCLs, also known as K-levels (K0 through K4), are [19]: Level 0: Does not have the ability or potential to ambulate or transfer safely with or without assistance and a prosthesis does not enhance their quality of life or mobility. Level 1: Has the ability or potential to use a prosthesis for transfers or ambulation on level surfaces at fixed cadence. Typical of the limited and unlimited household ambulator. Level 2: Has the ability or potential for ambulation with the ability to traverse low-level environmental barriers such as curbs, stairs, or uneven surfaces. Typical of the limited community ambulator. Level 3: Has the ability or potential for ambulation with variable cadence. Typical of the community ambulator who can traverse most environmental barriers and may have vocational, therapeutic, or exercise activity that demands prosthetic utilization beyond simple locomotion. Level 4: Has the ability or potential for prosthetic ambulation that exceeds basic ambulation skills, exhibiting high impact, stress, or energy levels. Typical of the prosthetic demands of the child, active adult, or athlete.

Across ambulatory individuals with lower limb amputation seeking prosthetic services, 8% are classified as K1, 30% as K2, 48% as K3, and 14% as K4 [8]. For those who are prescribed a prosthesis, a summary of available evidence reveals 24% 29% are only using it indoors at their 1-year follow-up, suggesting some may not be reaching their potential MFCL [20]. Among US service personnel who experienced lower limb amputation from participating in a military conflict in the last two decades, 6% were household ambulators (akin to K1), 13% were community ambulators (K2), 23% could walk with varying speed (K3), 26% participated in low-impact activities (also K3), and 26% in high-impact activities (K4) [21]. At the lower end of the function spectrum, prosthesis abandonment may be a concern. Balk et al.’s [20] summary of available evidence suggests 11% 22% have stopped using the prosthesis by their 1-year follow-up and that those with transfemoral amputations are about twice as likely to abandon as those with transtibial amputations. Among US service personnel at least 1-year postamputation, 10% were not currently using their prosthesis [21]. The ability to ambulate is central to the MFCLs K1 through K4 as it is fundamental to all daily activity involving personal transportation. Among US adults with a mobility disorder, walking is the most common physical activity (34%) followed by conditioning exercises (9%), lawn and garden related activities (7%), and household activities including childcare (3.2%) [22]. Despite the apparent simplicity, walking can often include additional challenges. In a North American study, individuals with lower limb amputation were queried on the ability to walk in different circumstances [23]. Among those with transtibial (transfemoral) amputations who could walk unassisted, 88% (87%) could do so in their home, 76% (73%) outside on level ground, 63% (47%) while carrying an object, 48% (46%) outside on uneven ground, and 48%

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(37%) down a few steps without a handrail. The majority could perform basic ambulatory activities (walk in or outside the house on level ground), but more advanced activities quickly became difficult.

48.2.1 Activity levels A direct measure of walking activity can be obtained using one of many, widely available wearable instruments that output steps taken per day. Leg-mounted instruments may be more accurate, and some can count very slow walking shuffling steps that would otherwise be missed [24]. This may be less important if the intent is counting steps that contribute to health which emphasize participating in moderate-to-vigorous activities rather than household activities of daily living. In reviewing the literature, it is important to note where the instrument is worn as leg-mounted instruments count steps per affected limb whereas waist-mounted instruments count both legs (i.e., twice as many steps). Normative data indicates adults approximately 20 65 years of age and living without disability or chronic illness take between 4000 and 18,000 steps/day [25]. Computed estimates for a reasonable threshold for free-living activity levels that include habitual daily activities plus 30 minutes of moderate-to-vigorous physical activities five times per week and some “off” days over a 7-day week, is 7100 11,000 steps/day [25]. Normative data on older adults ( . 65 years of age) who do not have a disability or chronic illness that limits their mobility or physical endurance, indicates they take between 2000 and 9000 steps/day [24]. Computed estimates of a reasonable threshold to include habitual daily activities (of about 5000 steps/day) and moderate-to-vigorous physical activities is 8000 steps/day (7100 steps/day if averaged over a 7-day week) [24]. Populations that have a disability or chronic illness that limits their mobility and or physical endurance take between 1200 and 8800 steps/day [24]. A computed estimate for a reasonable threshold for free-living activity levels is 5500 steps/day (4600 steps/day if averaged over a 7-day week) [24] and is lower because these individuals are likely to have a lower amount of habitual activity. Holden et al. [26] suggested individuals with a lower limb amputation who have a moderate level of support from family or social agencies (i.e., MFCL K1), must walk a minimum of 1200 steps/day to live in a single-level house or apartment and perform basic daily activities such as meal preparation, laundry, and light cleaning. To live independently in such an environment without support and performing some activities outside of the home (i.e., MFCL K2), would require between 2200 and 2900 steps/day [26]. In terms of the threshold for habitual daily activities and moderate-to-vigorous physical activity, some studies have shown individuals with lower limb amputation can exceed the target, 4600 5500 steps/day threshold, some fall within the threshold, and others are below it. (e.g., [27 29]). The wide variation exhibited in activity levels is to be expected based on the MFCL descriptions. While activity level measurements can provide a general idea of prosthetic foot use, they have not been able to reveal differences in prosthetic feet. Even comparisons between vastly different prosthetic feet failed to show a difference in step activity between devices [30 32]. It may be that comparative studies involving different feet will only reveal differences when steps are counted during specific activities like walking on ramps or on uneven terrain.

48.2.2 Activity bouts and durations Many wearable devices output steps/day, but shorter periods can provide additional information about how an individual with a lower limb amputation is using their prosthesis [33]. Counting steps over a 1-minute interval revealed that most activity bouts performed by individuals with lower limb amputation are less than a few minutes in duration. These short duration bouts of 1 2 minutes recorded less than 17 steps/minute. During medium duration bouts of around 10 minutes, which occurred only once or twice each day, the rate was much higher at approximately 70 steps/minute. Long duration bouts of 20 minutes or so, occurred only once every few days with a slower rate of approximately 50 steps/minute. These results suggest medium and long duration bouts are to a destination and may largely consist of steps taken in a straight line. The slower rate of long duration bouts, compared to medium duration bouts, may be due to the individual pacing themselves to avoid fatigue or injury. In contrast with medium and long duration bouts, short duration bouts with few steps suggest the occurrence of turning or maneuvering steps. Architectural barriers found in the household or within an office are likely to provide constraints.

48.2.3 Activity in different environments Video task analysis can provide a more detailed examination of ambulation context and environment. A preliminary exploration of the household environment with older adults ( . 65 years of age), who likely did not have a disability or chronic illness that limited their mobility or physical endurance (n 5 4), found turning steps comprised 13% 24% of

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all steps across eight different household activities [34]. One of every 5 steps was a turn, of which nearly 50% were turns between 76 and 120 degrees. A community-based video task analysis observed younger adults (n 5 11), who did not have a disability or chronic illness that limited their mobility or physical endurance, walking from one office to another (47 m total distance), through a cafeteria (68 m), through a convenience store (82 m), and from an office to a car in the parking lot (252 m) [35]. One of every 4 steps was a turn but differed by activity. Eight percent of steps were turns when walking from an office to the parking lot, compared to 35% turning steps in the convenience store, 45% turning steps when walking from one office to another, and 50% of all steps in a cafeteria. Ambulatory individuals with lower limb amputation can perform a variety of activities (e.g., walking straight, turning), but video-based methods observing specific activities in the household and community environments may constrain the results. Observing the step count distribution of different activities in free ranging individuals with lower limb amputation requires a higher fidelity instrument. Using sensors placed on or in the prosthetic pylon to record leg motions and a machine learning algorithm to categorize the signals, data was continuously collected over 1 2 days on individuals with transtibial amputation who were free to pursue their usual activities [36]. When walking about the home, work, and community environments, 82.8% of all steps were in a straight line, 9.0% were turning steps, 4.8% were steps on stairs, and 3.6% were steps on ramps. Turns to the left and right were equal, as were turns taken with the prosthesis on the inside or the outside of the turn. Steps taken while walking up and down ramps were also equal, as were steps taken walking up and down stairs. In summary, the activity levels of individuals with lower limb amputation are lower than those without mobility disabilities, can exhibit a large range but are mostly of short duration, and nearly 1 in 5 of all steps taken involved turning or walking on stairs and ramps.

48.3

Form of prosthetic feet

There are five traditional categories that are widely used by clinicians to describe prosthetic feet (e.g., [37]). The five categories include single axis, solid ankle cushion heel (SACH), flexible keel, multiaxis, and energy storing and returning (ESAR) prosthetic feet. Single-axis feet have a single hinge that allows plantar and dorsiflexion motion. This motion increases stability by allowing the sole of the foot to quickly come into full contact with the ground. Different elastomeric springs or bumpers can be installed to separately control plantar- and dorsiflexion motion at the discretion of the practitioner. SACH feet were developed in the late 1950s to incorporate the functions of the single-axis foot into an integrated design. The foot is nonarticulating but includes a compressible element (available in different stiffnesses) in the heel that mimics shock absorption and plantarflexion at initial ground contact and a rigid keel providing midstance stability at the expense of lateral movement. The reduced maintenance, cost, and weight of the SACH foot supplanted the single-axis foot. Flexible keel feet have a nonarticulating keel that can deform under load, but have limited capacity to return energy. Multiaxis feet provide plantarflexion and dorsiflexion, inversion and eversion, and internal and external rotation to accommodate a variety of terrains. Multiaxis feet can be of multipart designs, split keels, or a plate and urethane sandwich. ESAR (a.k.a., dynamic response) feet emerged in the 1980s and refer to feet that can store energy during stance and return it. To allow footwear, some ESAR feet require a separate foot cover or shell to be donned over the forefoot and heel keels, while other ESAR feet have an internal keel built into the foot comprising a single, combined unit. Some prosthetic feet belong to more than one of these traditional categories. The last two decades have seen significant innovations in prosthetic feet suggesting the need for new or additional categories described by their engineering features (e.g., [38]), but the traditional categories remain relevant due to the Medicare coverage guidance that describe the coverage determination of prosthetic feet [19]. The guidance states that prosthetic foot determination is made by the treating practitioner and/or prosthetist based upon the functional needs of the individual and there must be sufficient clinical documentation of functional need for a given type of foot. Basic lower limb prostheses should include a SACH foot but other prosthetic feet may be considered for coverage based upon the MFCL. The guide uses Medicare Healthcare Common Procedure Codes System (HCPCS) codes (known as L codes) that allow the prosthetist to assemble a prosthesis that is reasonable and necessary. Determinations of the type of prosthetic foot are summarized in the following: G G G

A SACH or single-axis foot is covered for MFCL K1 or above. A flexible-keel foot or multiaxial foot is covered for MFCL K2 or above. A microprocessor-controlled ankle-foot system, energy storing foot, dynamic response foot with multiaxial ankle, flex foot system, flex-walk system or equal, or shank foot system with vertical loading pylon is covered for MFCL K3 or above.

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G

753

Two types of prosthetic feet are specifically denied coverage as not reasonable and necessary. The microprocessor foot or ankle system addition with power assist which includes any type of motor and user-adjustable heel height feature.

The Centers for Medicare and Medicaid Services hold periodic hearings to consider and make coverage determinations for new innovations, but new L codes are an infrequent occurrence.

48.3.1 Solid ankle cushioned heel The SACH foot’s cushioned heel can be achieved with different designs. The 1950s design consisted of a cushioning wedge (see Fig. 48.1), available in three different stiffnesses, built into the heel [39]. Also available are cylinder shaped inserts of different stiffnesses to achieve the same result. If the compressible element is too soft, foot flat can occur prematurely, resulting in a foot that feels “slow” [37]. It can also delay the time from initial contact to loading response, making for asymmetric gait [40]. A too stiff compressible element defeats the purpose of the design.

48.3.2 Fixed-angle stiffness A fixed-angle stiffness prosthetic foot [38] has been described as a leaf spring whose foot-pylon angle becomes fixed after prosthesis assembly and alignment by a prosthetist. The forefoot keel of fixed-ankle stiffness prosthetic feet typically extends proximally into the ankle or shank and a cover is donned over the foot to accommodate footwear (see Fig. 48.2). Another design uses a flexible internal keel made from nylon, or similar material, constructed integral with

FIGURE 48.1 Cross-section of a solid-ankle cushion-heel (SACH) prosthetic foot (SACH; Kingsley Manufacturing Co., Costa Mesa, CA).

¨ ssur, Foothill Ranch, CA). Shown without FIGURE 48.2 Fixed-angle stiffness foot constructed with forefoot and heel keels (Vari-Flex with EVO; O footshell.

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FIGURE 48.3 Fixed-angle stiffness foot constructed with an internal keel integral to the foot (Seattle Lightfoot2; TruLife, Poulsbo, WA).

FIGURE 48.4 Fixed-angle stiffness feet constructed with a progressive or stiffening spring structure (Thrive; Freedom Innovations, Irvine, CA). Shown without footshell.

the foot (see Fig. 48.3). The stiffness of these feet can depend on cross-section, material, length, contour geometry, and sometimes the addition of multiple keels that collapse progressively (see Fig. 48.4) [37]. To accommodate uneven terrain, some incorporate a split keel. Many allow the prosthetist to preferentially stiffen the heel by adding elements between the heel and forefoot keels. Most manufacturers of fixed-angle stiffness prosthetic feet offer multiple stiffnesses within each product line that are selected based on the wearer’s body mass and self-reported activity level.

48.3.3 Variable-angle stiffness A variable-angle stiffness prosthetic foot is also a leaf spring but one whose foot-pylon angle can be varied after prosthesis assembly and alignment by a prosthetist. One design uses a microprocessor controlled, low power, nonbackdrivable motor to change the plantar- and dorsiflexion alignment angle during swing phase to increase toe clearance and adapt to changing slopes and ramps. Other designs use a damping element to control plantar- and dorsiflexion motion allowing the alignment angle to change during stance. Some have damping that are at fixed settings (see Fig. 48.5) and others use microprocessors and sensors to detect and respond to changes in terrain. Two particularly intriguing, but not commercially available, variable-angle stiffness feet include one with both a low power, nonbackdrivable motor and a damping element to make angle changes during both swing and stance [38,41] and another that uses a shear thickening fluid to create a variable damping element [42].

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FIGURE 48.5 Variable-angle stiffness foot that uses fixed setting damping to change alignment angle during stance (Echelon; Blatchford, Miamisburg, OH). Shown without footshell.

Prosthetic feet with microprocessors may provide additional benefits (e.g., enhanced ability to respond to changing environments) but rely on batteries to operate. Their additional weight, the need to recharge the batteries, susceptibility to water damage, and added cost may limit their appeal.

48.3.4 Variable stiffness While most of the previously discussed prosthetic feet are commercially available, variable stiffness feet have not yet reached the market. Commercialization may be on the horizon as various designs have been reported in the literature that vary plantarflexion and dorsiflexion stiffness (e.g., [43,44]), inversion and eversion stiffness [45], and transverse plane stiffness [46,47].

48.3.5 Powered prosthetic feet Powered prosthetic feet are an attractive research topic as they offer the ability to provide propulsion. Liu [48] conducted an extensive systematic review to classify and compare the mechanical designs of powered ankle-foot prostheses from 2000 to late 2019. Ninety-one powered ankle-foot prostheses were described and categorized primarily by actuator type (hydraulic, pneumatic, or electromechanical) and transmission mechanism (basic, multibar, or cam-based). Actuators with elastic elements were further classified by the assembly position of the elastic element. Most designs are based on a box rule where a subject of a specific mass walks at a specific speed which can be used to specify the actuator power requirements (typically between 250 and 350 W) [49]. The use of elastic elements in series, parallel, or in various combinations, can reduce the power requirement to between 50 and 200 W [49], resulting in substantial weight savings. Common control strategies include a high-level finite-state controller identifying terrain and gait cycle state with a low-level feedback control of actuator torque. Despite the 91 different devices reported in the last two decades, only one powered ankle-foot prosthesis is commercially available [50]. The BiOM T2 (BionX; Bedford, MA) was an earlier powered device and was marketed from 2012 to 2017. The Empower (Ottobock; Austin, TX) is the only currently available device and has been on the market since 2018 (see Fig. 48.6). Cost and coverage are significant impediments to more widespread use.

48.4

Function of prosthetic feet

To explore the function of prosthetic feet, investigators can perform human subject experiments, conduct mechanical property tests, and use musculoskeletal modeling and simulation. Human subject experiments can include questionnaires to discover user perception (e.g., Prosthesis Evaluation Questionnaire [51], Prosthetic Limb Users Survey of Mobility [52]), but most studies involve movement analysis. Many are performed on level ground, but stairs, ramps, and uneven terrains are also used. Speed is sometimes controlled as an independent variable, but self-selected is the most common. Spatiotemporal metrics such as stride length and cadence are widely reported. Joint kinematics and kinetics are also common, but ankle motion should be interpreted with caution [53]. Ground reaction forces and center

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FIGURE 48.6 Powered prosthetic ankle-foot (Empower; Otto Bock HealthCare LP, Austin, TX). Shown without footshell.

of pressure (COP) are also widely studied, as is whole-body angular momentum as a proxy for stability. Metabolic measurements are occasionally reported. More recent studies are using wearable devices to measure performance outside the laboratory. Human subject studies involving individuals wearing lower limb prostheses can be confounded by alignment, socket and suspension systems, residual limb length, and footwear. The length of acclimation period may also be a factor. Humans are extremely adaptive by nature which can make it hard to distinguish differences between feet. Mechanical properties testing resolves problems of human subject variability. Load versus deformation tests can yield estimates of prosthetic foot stiffness and loading-unloading curves can reveal hysteric losses. In conjunction with pseudo-static load-deformation tests, dynamic tests at specified velocities can be used to estimate damping. Impact tests can estimate shock absorption and natural frequency tests may have applications for the running prostheses of elite athletes. Musculoskeletal modeling and simulation offer tools to explore issues that are impractical or infeasible to conduct with human subjects. They can be used to discover prosthesis and muscle (both intact and residual limb) contributions to body weight support and propulsion, evaluate performance, and optimize prosthesis designs on a population or userspecific basis.

48.4.1 Clinical trials The designs of many clinical trials involving prosthetic feet measure their efficacy. The efficacy of a device, in comparison with another device such as a standard-of-care, is determined in randomized controlled trials (RCT) with controls governing the inclusion and exclusion of participants, environmental conditions, and other variables that might introduce bias or confound the results. Efficacy studies are usually performed in a laboratory environment. In contrast, some clinical trials measure effectiveness. The effectiveness of a device, also in comparison with another device, is determined in a pragmatic study that is much more reflective of the real-world conditions that would be encountered in the practice of medicine. Pragmatic studies have few inclusion and exclusion criteria or environmental controls. Effectiveness studies are typically performed in the individual’s usual environment (e.g., home, community, work). A few studies can be difficult to classify as there is a continuum between RCT and pragmatic studies [54]. Most clinical trials involving prosthetic feet have compared the efficacy of SACH versus ESAR feet [55]. These studies have used spatiotemporal, kinematics, kinetics, electromyography, and metabolic measures and the results provide strong support for the recommendation that patients capable of variable speed and/or community ambulation would benefit from ESAR feet prescriptions.

48.4.1.1 Clinical trials examining effects of stiffness The effects of prosthetic foot stiffness on biomechanical metrics have been measured in numerous clinical trials. Importantly, a 7.7% change in stiffness can be detected with 75% accuracy [56]. A summary of clinical trials examining the effects of stiffness [57] is found in the following: G G

More compliance results in greater range of motion ([43,57 63], and also [57]), More compliance results in greater power generation [64],

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G G G

757

More compliance results in increased prosthetic push-off peak power and work [65], More compliance results in greater energy return [60,62,64,66,67], More compliance results in increased spatial symmetry [68], and another study reported that

G

High stiffness is needed to maximize standing stability [69].

An important design trade-off exists between the need to support body weight with sufficient stiffness and the desire to maximize energy return via more compliance. Stark [37] suggested prosthetic foot designers would tend to err toward a stiffer forefoot to avoid a drop-off effect at terminal stance, an opinion later confirmed experimentally [58]. A few clinical trials have reported preferences related to prosthetic foot stiffness. After walking both overground and on a treadmill with five custom feet of differing forefoot stiffness, transtibial participants ranked the most compliant forefoot as the worst and preferred the middle stiffness and somewhat stiffer forefeet [70]. Outside of the laboratory, a comparison of stiff, intermediate, and compliant commercially available feet found the most compliant foot was preferred after two of the three feet were worn (presumably) in the household, community, and work environments for 1week [71]. In contrast, a study examining three consecutive stiffness categories of a commercially available foot found participants were inconsistent in ranking their order, suggesting the inability to perceive differences across a clinically relevant range [65]. To study the effect of walking speed on preferred stiffness, transtibial participants fit with a novel variable-stiffness prosthetic ankle-foot (i.e., varying not just forefoot stiffness), preferred the most compliant setting at their self-selected treadmill speed compared to faster and slower speeds [72]. When walking overground, these individuals walked faster with a stiffness at or above their preferred [72]. Surprisingly, their prosthetists preferred a stiffness 26% higher than they did [73].

48.4.1.2 Clinical trials examining effects of damping The effects of variable-angle stiffness achieved with fixed setting damping elements have also been measured in several clinical trials. A summary [74] found the following: G G G G

Socket pressure and loading rates on residual limb were reduced [75], Compensatory joint moments were reduced [76], Toe clearance was increased [77], and Self-selected walking speed was increased [76 79].

48.4.1.3 Powered prosthetic feet Clinical trials have compared the performance of powered prosthetic feet and ESAR prosthetic feet on level ground, slopes, and uneven terrain, finding that: G

G

G

G

G

G

Metabolic energy expenditure effects on level ground are inconclusive, with some studies showing powered prostheses reduce metabolic costs [80,81] and others no significant differences [82 84]. Preferred walking speed effects are inconclusive. On level ground some studies showing an increase while wearing a powered prosthesis [80,85] and others no significant effects [84]. On a loose rock surface, preferred walking speed was shown to increase while wearing a powered prosthesis [86]. Stability metrics are also inconclusive. On level ground, subjects wearing the powered prosthesis have shown a decrease in the sagittal plane range of whole-body angular momentum but not in the coronal plane [87]. On uneven terrain, a decrease in center of mass medial-lateral motion was observed with the powered prosthesis, but no commensurate change in the medial-lateral margin of stability [86]. Biomechanic metrics related to injuries to the unaffected limb have shown reduced peak forces and knee adduction moment when wearing a powered prosthesis, but the effects were not uniform across speeds of 0.75, 1.00, 1.25, 1.50, and 1.75 m/s [88]. Perception of mobility with a powered prosthesis did not correlate with daily activity levels. Subjects self-reported increased ambulation, but daily step counts did not change [84]. Perception of social burden and frustration was decreased with a powered prosthesis, but there were no differences in quality of life or a preference for the powered device [84]. Tuning of the powered prosthesis may play an important role in the performance of the device.

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Metabolic energy expenditure on a slope (15 degrees incline) showed no significant difference [81] when the prosthesis was tuned by a manufacturer’s representative to patient preference within a normative range, presumably during level ground walking. However, tuning the powered prosthesis on specific slopes to reach equivalence in ankle range of motion, peak moment, power, and net work of the unaffected limb and/or within 2 standard deviations of the mean of individuals without lower limb amputation, yielded significant differences [83]. Walking at 1.25 m/s on slopes of 0, 6 3, 6 6, and 6 9 degrees with the powered prosthesis tuned at each slope (an arduous and nonclinical task), revealed net metabolic power reductions of 5% at 13 and 16 degrees. Results from 19 suggested an even greater reduction of 13% but was not statistically significant as only 2 of 10 subjects were able to complete the protocol. Metabolic cost of transport on level ground can be influenced by tuning [89]. Six different power settings (0%, 25%, 50%, 75%, 100%, and a prosthetist-chosen setting) were compared while participants walked on a treadmill. Power settings of 0% and 25% had higher cost of transport compared to settings of 50%, 75%, and 100%. Interestingly, there was always a higher power setting, but not the maximum, that would result in lower cost of transport than the prosthetist-chosen setting.

The ability of the individual to adapt and learn how to best use a powered prosthesis may also play a role in its outcomes as [82,84,89] all noted their results varied considerably between subjects. A secondary analysis in [82] examined various user characteristics to see if they influenced cost of transport. Age, residual limb length, body mass, years of prosthesis experience, and strength had no effect. Only MFCL, specifically levels K3 and K4, had a relationship with the outcome. K4s were more likely to exhibit decreased metabolic cost than K3s. To understand this unexpected result an exploratory analysis was conducted and found decreased metabolic cost was moderately correlated with increases in residual limb gluteus medius, quadriceps, and hamstrings muscle activity after selective data exclusions [90]. This was another unexpected result, namely energy savings occurred with more muscle activity. The quadriceps were mostly active at the end of stance, which the authors suggested may be a strategy to stiffen the knee joint to better use the propulsive power of the prosthesis. The hamstrings were mostly active during swing, which the authors suggested may be a strategy to reduce excessive prosthetic leg swing. K4 ambulators would likely be able to execute these strategies while K3 ambulators might not. Not all individuals who are prescribed a powered prosthesis continue to use it. A survey of individuals who had been fit with a powered prosthesis found 61% had abandoned the technology and reverted to a passive prosthesis [50]. Most of those who continued with the powered prosthesis believed its use reduced their pain and improved their mobility. One reason for abandonment is that learning to effectively use the device may not be intuitive, suggesting a gait training program oriented toward powered prostheses might help to increase the number of individuals who could benefit from its prescription.

48.4.2 Mechanical property tests Mechanical property tests offer an objective comparison between prosthetic feet tests. The International Organization for Standardization (ISO) has published standards for the structural testing of lower limb prostheses (ISO 10328:2016, [91]) and the testing of ankle-foot devices and foot units (ISO 22675:2016 [92]). Additional standardized tests to validate specific features related to Medicare HCPCS L codes for prosthetic feet have also been developed by the American Orthotic and Prosthetic Association and participating manufacturers and was published as a white paper [93]. These have been adopted by some researchers [44,94,95].

48.4.2.1 Stiffness, hysteresis, and energy A comprehensive load versus deformation test in 1990 on 9 different commercially available prosthetic feet set the standard for mechanical property tests [96]. Van Jaarsveld varied the angle between the pylon from 2 30 degrees (heel contact) to 135 degrees (toe-off) in 1 degrees increments for each test (66 in total) which loaded the foot to 1000 N or 35 mm deformation (whichever occurred first). The mean stiffness was calculated at maximum load or deformation, and hysteresis was calculated as the area between the loading and unloading curves, allowing objective comparison between the different feet. A more recent test on 7 different, commercially available ESAR prosthetic feet used ISO 10328 [91] as a guide to obtain objective measures [97]. Tests were performed at a pylon angle of 2 15 degrees (heel region) and 120 degrees (toe region) with the foot aligned at 7 degrees toe out. The feet were loaded to 2240 N per the standard, but stiffness was calculated over a 6 5% range at a walking load of 1068 N and at a moderately paced running load of 1780 N.

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Hysteresis was calculated based on the maximum load and expressed as a percentage of both input energy and absolute energy loss. Clear differences were readily observable between feet. To discover the stiffnesses that might be experienced when walking on varying terrain or when turning, loaddeformation tests have also been conducted in the sagittal plane [96,97], but also in the coronal plane [98]. Seven different feet, including a range of stiffnesses for each manufacturer and the effects of a heel stiffening wedge in two of the feet, with the angle of the pylon in 15 sagittal plane pylon positions and 5 coronal plane pylon positions for each test (75 in total). The load was based on scaled vertical ground reaction force (vGRF) data as a function of pylon angle for walking at 1 m/s. Stiffness was calculated over a 6 1 standard deviation range of the vGRF and energy storage was calculated by integrating the load-deformation over the load range. The resulting stiffness and energy storage plots (a link graphical user interface allowing individuals to make their own three-dimensional comparisons is provided in [98]) allow an objective comparison for a variety of potential, everyday ambulatory conditions.

48.4.2.2 Roll-over shape A method for analyzing the function of prosthetic ankle-feet results in a roll-over shape and is derived from the measured COP from the net force on a prosthetic ankle-foot under loads that simulate walking (or walking itself) and transform them to a local coordinate system at the attachment point of a pylon or socket [99,100]. The COP path in the local coordinate system reveals the effective rocker shape, or roll-over shape, taken by the prosthetic foot between heel contact and opposite heel contact. Local coordinates whose vertical axis is aligned with the shank (using ankle and knee reference points) produce ankle-foot roll-over shapes and may be more relevant to understanding the function of transtibial prostheses. Local coordinates whose vertical axis is aligned with the leg (using ankle and hip reference points) produce knee-ankle-foot roll-over shapes and may be most relevant to understanding the function of transfemoral prostheses. Studies on individuals without lower limb amputations have shown the roll-over shape is invariant to walking speed [101], added weight [102], and heel height [103]. The different mechanical properties of prosthetic ankle-feet can exhibit distinctly different roll-over shapes, but experienced prosthetists tend to align them such that their roll-over shapes overlap [104] using slide, tilt, or various adapter combinations [37]. The finding that distinctly different prosthetic ankle-foot roll-over shapes, such as those produced from prosthetic feet of different stiffness categories from the same manufacturer, become similar after prosthetist alignment may explain the null results of many clinical trials.

48.4.3 Musculoskeletal modeling and simulation Musculoskeletal modeling and simulation have been widely used to analyze human movement. Models of individuals with lower limb amputations have been used to gain insight into individual muscle and prosthetic ankle-foot contributions to walking mechanics [105,106], intersegmental knee loading across steady-state walking speeds [107], knee joint contact forces (which cannot be directly measured) during walking [108], the influence of ESAR properties on leg loading [59], and the relationships between prosthetic ankle-foot stiffness and muscle function during walking [109]. Models distributing the stiffness of the prosthetic ankle-foot have been used to generate forward dynamics walking simulations for identifying the optimal stiffness profile that best emulated experimental walking mechanics while minimizing metabolic energy expenditure and intact leg internal knee contact forces [110]. Additionally, musculoskeletal modeling and simulation has also been used gage quadriceps and hamstrings strength needed to maintain metabolic costs after amputation [111]. These studies have provided clinically meaningful insight into individual muscle and prosthesis contributions to walking performance.

48.5

Future prosthetic foot research

New prosthetic feet are introduced to the market each year, but little objective information is available to appreciate their place in prescription practice. Adoption and reporting of standardized mechanical property tests (e.g., [93]) would enable objective comparisons with existing products [112]. Wearable technologies capable of measuring prosthetic foot use and context [e.g., activity level, duration, intensity, location (household or community), terrain (ramps, stairs, uneven ground, slippery conditions)] in conjunction with electronic health record information (e.g., prosthetic prescription, residual limb injuries, injurious falls) would enable

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identification of benefits (e.g., improved mobility) and risks (e.g., incidence of injury) associated with specific prosthetic prescriptions for particular patient characteristics or populations. The differences between SACH and ESAR prosthetic feet are well-established, but evidence for other prosthetic feet is needed [16]. Clinical trials, mechanical property tests, and musculoskeletal modeling and simulation studies to determine differences between articulated and nonarticulated prosthetic feet would aid in clinical practice. An improved understanding of why some individuals benefit from powered prostheses and others do not is needed. Development of powered prosthesis gait training and strength programs might help to increase the number of individuals who can benefit from this technology [50]. Existing prosthetic feet, both commercially available and research prototypes, are no match for the natural limb. Development of prosthetic feet that can perform task-specific activities (e.g., conform to terrain, drive a vehicle, enhance balance) would be useful to individuals with lower limb amputations. Finally, current lower limb prostheses are devoid of systems that allow the individual to sense and respond to their environment. Sensory feedback would be a meaningful improvement.

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[50] Kannenberg A, Morris AR, Hibler KD. Free-living user perspectives on musculoskeletal pain and patient-reported mobility with passive and powered prosthetic ankle-foot components: a pragmatic, exploratory cross-sectional study. Front Rehabil Sci 2022;2:1 12. Available from: https://doi.org/10.3389/fresc.2021.805151. [51] Legro MW, Reiber GD, Smith DG, Del Aguila M, Larsen J, Boone D. Prosthesis evaluation questionnaire for persons with lower limb amputations: assessing prosthesis-related quality of life. Arch Phys Med Rehabil 1998;79:931 8. Available from: https://doi.org/10.1016/S0003-9993(98)90090-9. [52] Hafner BJ, Gaunaurd IA, Morgan SJ, Amtmann D, Salem R, Gailey RS. Construct validity of the prosthetic limb users survey of mobility (PLUS-M) in adults with lower limb amputation. Arch Phys Med Rehabil 2017;98:277 85. Available from: https://doi.org/10.1016/j. apmr.2016.07.026. [53] Zelik KE, Honert EC. Ankle and foot power in gait analysis: implications for science, technology and clinical assessment. J Biomech 2018;75:1 12. Available from: https://doi.org/10.1016/j.jbiomech.2018.04.017. [54] Patsopoulos NA. A pragmatic view on pragmatic trials. Dialogues Clin Neurosci 2011;13:217 24. [55] Stevens PM, Rheinstein J, Wurdeman SR. Prosthetic foot selection for individuals with lower-limb amputation: a clinical practice guideline. J Prosthet Orthot 2018;30:175 80. Available from: https://doi.org/10.1097/JPO.0000000000000181. [56] Shepherd MK, Azocar AF, Major MJ, Rouse EJ. Amputee perception of prosthetic ankle stiffness during locomotion. J Neuroeng Rehabil 2018;15:1 10. Available from: https://doi.org/10.1186/s12984-018-0432-5. ´ rmannsdo´ttir AL, Lecomte C, Brynjo´lfsson S, Briem K. Task dependent changes in mechanical and biomechanical measures result from [57] A manipulating stiffness settings in a prosthetic foot. Clin Biomech 2021;89. Available from: https://doi.org/10.1016/j.clinbiomech.2021.105476. [58] Klodd E, Hansen A, Fatone S, Edwards M. Effects of prosthetic foot forefoot flexibility on gait of unilateral transtibial prosthesis users. J Rehabil Res Dev 2010;. Available from: https://doi.org/10.1682/JRRD.2009.10.0166. [59] Ventura JD, Klute GK, Neptune RR. The effect of prosthetic ankle energy storage and return properties on muscle activity in below-knee amputee walking. Gait Posture 2011;33:220 6. Available from: https://doi.org/10.1016/j.gaitpost.2010.11.009. [60] Fey NP, Klute GK, Neptune RR. The influence of energy storage and return foot stiffness on walking mechanics and muscle activity in belowknee amputees. Clin Biomech 2011;. Available from: https://doi.org/10.1016/j.clinbiomech.2011.06.007. [61] Major MJ, Twiste M, Kenney LPJ, Howard D. The effects of prosthetic ankle stiffness on ankle and knee kinematics, prosthetic limb loading, and net metabolic cost of trans-tibial amputee gait. Clin Biomech 2014;29:98 104. Available from: https://doi.org/10.1016/j. clinbiomech.2013.10.012. [62] Adamczyk PG, Roland M, Hahn ME. Sensitivity of biomechanical outcomes to independent variations of hindfoot and forefoot stiffness in foot prostheses. Hum Mov Sci 2017;54:154 71. Available from: https://doi.org/10.1016/j.humov.2017.04.005. [63] Shell CE, Segal AD, Klute GK, Neptune RR. The effects of prosthetic foot stiffness on transtibial amputee walking mechanics and balance control during turning. Clin Biomech 2017;49:56 63. Available from: https://doi.org/10.1016/j.clinbiomech.2017.08.003. [64] Glanzer EM, Adamczyk PG. Design and validation of a semi-active variable stiffness foot prosthesis. IEEE Trans Neural Syst Rehabil Eng 2018;26:2351 9. Available from: https://doi.org/10.1109/TNSRE.2018.2877962. [65] Halsne EG, Czerniecki JM, Shofer JB, Morgenroth DC. The effect of prosthetic foot stiffness on foot-ankle biomechanics and relative foot stiffness perception in people with transtibial amputation. Clin Biomech 2020;80. Available from: https://doi.org/10.1016/j. clinbiomech.2020.105141. [66] Zelik KE, Collins SH, Adamczyk PG, Segal AD, Klute GK, Morgenroth DC, et al. Systematic variation of prosthetic foot spring affects centerof-mass mechanics and metabolic cost during walking. IEEE Trans Neural Syst Rehabil Eng 2011;19:411 19. Available from: https://doi.org/ 10.1109/TNSRE.2011.2159018. [67] Koehler-McNicholas S, Nickel E, Barrons K, Blaharski K, Dellamano C, Ray S, et al. Mechanical and dynamic characterization of prosthetic feet for high activity users during weighted and unweighted walking. PLoS One 2018;13:e0202884. [68] Major MJ, Twiste M, Kenney LPJ, Howard D. The effects of prosthetic ankle stiffness on stability of gait in people with transtibial amputation. J Rehabil Res Dev 2016;53:839 52. Available from: https://doi.org/10.1682/JRRD.2015.08.0148. [69] Koehler-McNicholas SR, Savvas Slater BC, Koester K, Nickel EA, Ferguson JE, Hansen AH. Bimodal ankle-foot prosthesis for enhanced standing stability. PLoS One 2018;13:1 18. Available from: https://doi.org/10.1371/journal.pone.0204512. [70] Klodd E, Hansen A, Fatone S, Edwards M. Effects of prosthetic foot forefoot flexibility on oxygen cost and subjective preference rankings of unilateral transtibial prosthesis users. J Rehabil Res Dev 2010;47:543 52. Available from: https://doi.org/10.1682/JRRD.2010.01.0003. [71] Raschke SU, Orendurff MS, Mattie JL, Kenyon DEA, Jones OY, Moe D, et al. Biomechanical characteristics, patient preference and activity level with different prosthetic feet: a randomized double blind trial with laboratory and community testing. J Biomech 2015;48:146 52. Available from: https://doi.org/10.1016/j.jbiomech.2014.10.002. [72] Clites TR, Shepherd MK, Ingraham KA, Wontorcik L, Rouse EJ. Understanding patient preference in prosthetic ankle stiffness. J Neuroeng Rehabil 2021;18:1 17. Available from: https://doi.org/10.1186/s12984-021-00916-1. [73] Shepherd MK, Rouse EJ. Comparing preference of ankle foot stiffness in below-knee amputees and prosthetists. Sci Rep 2020;10. Available from: https://doi.org/10.1038/s41598-020-72131-2. [74] Safaeepour Z, Eshraghi A, Geil M. The effect of damping in prosthetic ankle and knee joints on the biomechanical outcomes: a literature review. Prosthet Orthot Int 2017;41:336 44. Available from: https://doi.org/10.1177/0309364616677651. [75] Portnoy S, Kristal A, Gefen A, Siev-Ner I. Outdoor dynamic subject-specific evaluation of internal stresses in the residual limb: hydraulic energy-stored prosthetic foot compared to conventional energy-stored prosthetic feet. Gait Posture 2012;35:121 5. Available from: https://doi. org/10.1016/j.gaitpost.2011.08.021.

Prosthetic Feet Chapter | 48

763

[76] De Asha AR, Munjal R, Kulkarni J, Buckley JG. Walking speed related joint kinetic alterations in trans-tibial amputees: impact of hydraulic “ankle” damping. J Neuroeng Rehabil 2013;10. Available from: https://doi.org/10.1186/1743-0003-10-107. [77] Johnson L, De Asha AR, Munjal R, Kulkarni J, Buckley JG. Toe clearance when walking in people with unilateral transtibial amputation: effects of passive hydraulic ankle. J Rehabil Res Dev 2014;51:429 38. [78] De Asha AR, Munjal R, Kulkarni J, Buckley JG. Impact on the biomechanics of overground gait of using an “Echelon” hydraulic ankle-foot device in unilateral trans-tibial and trans-femoral amputees. Clin Biomech 2014;29:728 34. Available from: https://doi.org/10.1016/j. clinbiomech.2014.06.009. [79] De Asha AR, Johnson L, Munjal R, Kulkarni J, Buckley JG. Attenuation of centre-of-pressure trajectory fluctuations under the prosthetic foot when using an articulating hydraulic ankle attachment compared to fixed attachment. Clin Biomech 2013;28:218 24. Available from: https:// doi.org/10.1016/j.clinbiomech.2012.11.013. [80] Herr HM, Grabowski AM. Bionic ankle-foot prosthesis normalizes walking gait for persons with leg amputation. Proc R Soc B Biol Sci 2012;279:457 64. Available from: https://doi.org/10.1098/rspb.2011.1194. [81] Esposito ER, Whitehead JMA, Wilken JM. Step-to-step transition work during level and inclined walking using passive and powered anklefoot prostheses. Prosthet Orthot Int 2016;40:311 19. Available from: https://doi.org/10.1177/0309364614564021. [82] Gardinier ES, Kelly BM, Wensman J, Gates DH. A controlled clinical trial of a clinically-tuned powered ankle prosthesis in people with transtibial amputation. Clin Rehabil 2018;32:319 29. Available from: https://doi.org/10.1177/0269215517723054. [83] Montgomery JR, Grabowski AM. Use of a powered ankle foot prosthesis reduces the metabolic cost of uphill walking and improves leg work symmetry in people with transtibial amputations. J R Soc Interface 2018;15. Available from: https://doi.org/10.1098/rsif.2018.0442. [84] Kim J, Wensman J, Colabianchi N, Gates DH. The influence of powered prostheses on user perspectives, metabolics, and activity: a randomized crossover trial. J Neuroeng Rehabil 2021;18:1 13. Available from: https://doi.org/10.1186/s12984-021-00842-2. [85] Ferris AE, Aldridge JM, Ra´bago CA, Wilken JM. Evaluation of a powered ankle-foot prosthetic system during walking. Arch Phys Med Rehabil 2012;93:1911 18. Available from: https://doi.org/10.1016/j.apmr.2012.06.009. [86] Gates DH, Aldridge JM, Wilken JM. Kinematic comparison of walking on uneven ground using powered and unpowered prostheses. Clin Biomech 2013;28:467 72. Available from: https://doi.org/10.1016/j.clinbiomech.2013.03.005. [87] D’Andrea S, Wilhelm N, Silverman AK, Grabowski AM. Does use of a powered ankle-foot prosthesis restore whole-body angular momentum during walking at different speeds? Clin Orthop Relat Res 2014;472:3044 54. Available from: https://doi.org/10.1007/s11999-014-3647-1. [88] Grabowski AM, D’Andrea S. Effects of a powered ankle-foot prosthesis on kinetic loading of the unaffected leg during level-ground walking. J Neuroeng Rehabil 2013;10. Available from: https://doi.org/10.1186/1743-0003-10-49. [89] Ingraham KA, Choi H, Gardinier ES, Remy CD, Gates DH. Choosing appropriate prosthetic ankle work to reduce the metabolic cost of individuals with transtibial amputation. Sci Rep 2018;8:1 12. Available from: https://doi.org/10.1038/s41598-018-33569-7. [90] Kim J, Gardinier ES, Vempala V, Gates DH. The effect of powered ankle prostheses on muscle activity during walking. J Biomech 2021;124:110573. Available from: https://doi.org/10.1016/j.jbiomech.2021.110573. [91] International Organization for Standardization Technical Committee 168 Prosthetics and Orthotics. ISO 10328. Prosthetics—structural testing of lower-limb prostheses—requirements and test methods 140. ,https://www.iso.org/standard/70205.html.; 2016. [92] International Organization for Standardization Technical Committee 168 Prosthetics and Orthotics. ISO 22675. Prosthetics—testing of anklefoot devices and foot units—requirements and test methods. ,https://www.iso.org/standard/70203.html.; 2016: 91. [93] American Orthotic & Prosthetic Association. AOPA’s Prosthetic Foot Project. ,http://www.aopanet.org/wp-content/uploads/2013/12/ Prosthetic_Foot_Project.pdf.; 2010. [94] Warder HH, Fairley JK, Coutts J, Glisson RR, Gall K. Examining the viability of carbon fiber reinforced three-dimensionally printed prosthetic feet created by composite filament fabrication. Prosthet Orthot Int 2018;42:644 51. Available from: https://doi.org/10.1177/ 0309364618785726. [95] Major MJ, Scham J, Orendurff M. The effects of common footwear on stance-phase mechanical properties of the prosthetic foot-shoe system. Prosthet Orthot Int 2018;42:198 207. Available from: https://doi.org/10.1177/0309364617706749. [96] Van Jaarsveld HWL, Grootenboer HJ, Koopman HFJM, de Vries J. Stiffness and hysteresis properties of some prosthetic feet. Prosthet Orthot Int 1990;14:117 24. Available from: https://doi.org/10.3109/03093649009080337. [97] Webber CM, Kaufman K. Instantaneous stiffness and hysteresis of dynamic elastic response prosthetic feet. Prosthet Orthot Int 2017;41:463 8. Available from: https://doi.org/10.1177/0309364616683980. [98] Womac N, Neptune R, Klute G. Stiffness and energy storage characteristics of energy storage and return prosthetic feet. Prosthet Orthot Int Appear 2019;. [99] Hansen AH. Scientific methods to determine functional performance of prosthetic ankle-foot systems. J Prosthet Orthot 2005;17:S23 9. [100] Hansen AH, Childress DS. Investigations of roll-over shape: implications for design, alignment, and evaluation of ankle-foot prostheses and orthoses. Disabil Rehabil 2010;32:2201 9. Available from: https://doi.org/10.3109/09638288.2010.502586. [101] Hansen AH, Childress DS, Knox EH. Roll-over shapes of human locomotor systems: effects of walking speed. Clin Biomech 2004;19:407 14. Available from: https://doi.org/10.1016/j.clinbiomech.2003.12.001. [102] Hansen AH, Childress DS. Effects of adding weight to the torso on roll-over characteristics of walking. J Rehabil Res Dev 2005;42:381 90. Available from: https://doi.org/10.1682/JRRD.2004.04.0048. [103] Hansen AH, Childress DS. Effects of shoe heel height on biologic rollover characteristics during walking. J Rehabil Res Dev 2004;41:547 53. Available from: https://doi.org/10.1682/JRRD.2003.06.0098.

764

PART | 7 Clincial Interventions

[104] Hansen AH, Meier MR, Sam M, Childress DS, Edwards ML. Alignment of trans-tibial prostheses based on roll-over shape principles. Prosthet Orthot Int 2003;27:89 99. Available from: https://doi.org/10.1080/03093640308726664. [105] Silverman AK, Neptune RR. Muscle and prosthesis contributions to amputee walking mechanics: a modeling study. J Biomech 2012;45:2271 8. Available from: https://doi.org/10.1016/j.jbiomech.2012.06.008. [106] Zmitrewicz RJ, Neptune RR, Sasaki K. Mechanical energetic contributions from individual muscles and elastic prosthetic feet during symmetric unilateral transtibial amputee walking: a theoretical study. J Biomech 2007;40:1824 31. Available from: https://doi.org/10.1016/j. jbiomech.2006.07.009. [107] Fey NP, Neptune RR. 3D intersegmental knee loading in below-knee amputees across steady-state walking speeds. Clin Biomech 2012;. Available from: https://doi.org/10.1016/j.clinbiomech.2011.10.017. [108] Silverman AK, Neptune RR. Three-dimensional knee joint contact forces during walking in unilateral transtibial amputees. J Biomech 2014;47:2556 62. Available from: https://doi.org/10.1016/j.jbiomech.2014.06.006. [109] Fey NP, Klute GK, Neptune RR. Altering prosthetic foot stiffness influences foot and muscle function during below-knee amputee walking: a modeling and simulation analysis. J Biomech 2013;46:637 44. Available from: https://doi.org/10.1016/j.jbiomech.2012.11.051. [110] Fey NP, Klute GK, Neptune RR. Optimization of prosthetic foot stiffness to reduce metabolic cost and intact knee loading during below-knee amputee walking: a theoretical study. J Biomech Eng 2012;134:111005. Available from: https://doi.org/10.1115/1.4007824. [111] Russell Esposito E, Miller RH. Maintenance of muscle strength retains a normal metabolic cost in simulated walking after transtibial limb loss. PLoS One 2018;13:e0191310. Available from: https://doi.org/10.1371/journal.pone.0191310. [112] Czerniecki JM. Research and clinical selection of foot-ankle systems. J Prosthet Orthot 2005;17:S35 7. Available from: https://doi.org/ 10.1097/00008526-200510001-00012.

Index Note: Page numbers followed by “f” and “t” refer to figures and tables, respectively.

A Abductor digiti minimi (ABDM), 33, 103 Abductor hallucis (ABH), 33, 103 Accessory navicular, 480 482, 707 Achilles tendons, 103, 313 314, 515 rupture, 474 thickness assessment in relation to weightbearing activities, 327 ultrasound combined with dynamometry to assess mechanical properties of, 329 331 elastography to assess mechanical properties, 333 334 Achilles tightness, 680 681 ACL-ITCL. See Anterior capsular ligament and interosseous talocalcaneal ligament (ACL-ITCL) Acoustic impedance, 324 Acquired flatfoot, 705 ACSA. See Anatomical cross-sectional areas (ACSA) Activity bouts and durations, 751 Activity in different environments, 751 752 Activity levels, 751 Actuation and assistance, 236 Acute ankle sprains, 473 Acute joint injury severity, 397 399 Adduction-abduction, 67 Adductor hallucis (ADH), 36 37, 103 Adipose tissue, 135 Adult-acquired flatfoot, reconstructions for failure of posterior tibial tendon, 679 680 forefoot external rotation, 674 677 gastrocnemius and Achilles’ tightness, 680 681 clinical assessment, 681 surgical treatment, 681 hindfoot valgus, 670 673 medial arch eversion, 681 682 bony anatomy, 682 clinical assessment, 681 682 ligamentous failure, 682 surgical reconstruction, 682 sag at talonavicular joint, 674 675 Adult-acquired flatfoot deformity (AAFD), 669 Advanced glycation end-products (AGEs), 420 AFO. See Ankle-foot orthosis (AFO) Aging, 108 109 areas for future research, 604

changing properties and functions of foot tissues, 595 599 foot function, 601 603 foot posture foot disorders, and mobility limitations, 603 604 and morphology, 599 601 AHI. See Arch height index (AHI) ALARA. See As low as reasonably achievable (ALARA) Allograft wedges, 676 677 American Society for Testing and Materials (ASTM), 127 Amputation level, 749 Analog radiography, 265 266 Anatomical cross-sectional areas (ACSA), 630 Anatomical descriptions, 54 64 application specific human anatomical positions, 56 61 foot motions, 64 71 persistence of eponyms, 54 regional descriptions, 54 55 Anatomical directions, planes, and axes, 61 64 foot midline, 64 posterior anterior vs. ventral dorsal, 61 Anatomical planes, 61 63 Anatomical position, 55 56 Anatomically-based coordinate systems, 153 Anatomy of foot, 3 blood supply, 41 44 muscles and fascial specializations, 20 38 nerves, 38 41 skeletal structures, 3 11 sequences, 303 Anisotropic materials, 414 415 Ankle, 77 78, 289, 311 313 alignment/malalignment, 738 ankle-high devices, 664 arthritis, 647, 651 652, 655 constitutive parameter identification for ankle ligaments, 420 dorsiflexion, 692 fusions, 698, 731 732, 739 740 nails, 727 injuries, 422 423, 461 instability, 510 joint, 11 13, 122 123 complex, 78 LCL, 122

MCL, 123 synovitis, 590 kinematics, 173 mortise width, 171 motion in three cardinal planes with arthroplasty, 737 movements, 422 423 replacement, 698 rocker, 649 Ankle and foot, 603 deformities, 438 related considerations and insights, 219 224 history-dependence of force, 221 224 maximal voluntary contractions and knee angle, 221 motor unit behavior and quantity, 219 220 Ankle arthrodesis, 732 733 areas of future biomechanical research, 744 biomechanical considerations/complications of, 738 740 ankle alignment/malalignment, 738 arthritis at distal joints, 738 739 cadaveric gait simulation of arthroplasty and arthrodesis, 739 740 component wear or failure, 739 computational models of arthroplasty and arthrodesis, 740 gait mechanics, 738 biomechanical factors in presurgical assessment and consideration of, 736 738 ankle motion in three cardinal planes with arthroplasty, 737 bony and ligamentous ankle anatomy with arthroplasty, 737 bony morphology and alignment with arthrodesis, 738 limb alignment with arthroplasty, 736 737 biomechanical outcomes, 741 brief description and history of surgical techniques, 732 735 clinical outcomes, 741 743 costs, 743 effectiveness, 742 743 patient subgroups, 743 safety, 741 742 Ankle arthroplasty, 733 735 areas of future biomechanical research, 744

765

766

Index

Ankle arthroplasty (Continued) biomechanical considerations/complications of, 738 740 ankle alignment/malalignment, 738 arthritis at distal joints, 738 739 cadaveric gait simulation of arthroplasty and arthrodesis, 739 740 component wear or failure, 739 computational models of arthroplasty and arthrodesis, 740 gait mechanics, 738 biomechanical factors in presurgical assessment and consideration of, 736 738 ankle motion in three cardinal planes with arthroplasty, 737 bony and ligamentous ankle anatomy with arthroplasty, 737 bony morphology and alignment with arthrodesis, 738 limb alignment with arthroplasty, 736 737 biomechanical outcomes, 741 brief description and history of surgical techniques, 732 735 clinical outcomes, 741 743 costs, 743 effectiveness, 742 743 patient subgroups, 743 safety, 741 742 Ankle osteoarthritis, 556 558. See also Midfoot osteoarthritis clinical and biomechanical effects of conservative treatment, 557 558 clinical findings, 556 557 etiology and impact, 556 structural and biomechanical features, 557 Ankle-foot orthosis (AFO), 344, 492, 647 650 altering line of action of ground reaction force, 648 controlling axial forces, 648 controlling rotational motion, 648 controlling translational motion, 648 design and prescription of, 652 654 articulated vs. nonarticulated, 653 654 conventional vs. advanced, 652 653 rocker bottom shoes, 654 655 materials, 655 new designs, 655 patient populations, 649 652 ankle arthritis, 651 652 cerebral palsy, 651 limb salvage, 652 stroke, 650 651 roll-over shape, 649 sport applications, 655 656 variations on materials, 655 Ankylosing spondylitis (AS), 589 Anterior capsular ligament and interosseous talocalcaneal ligament (ACL-ITCL), 124 Anterior impingement, 712 Anterior inferior tibiofibular ligament (AITFL), 311

Anterior talofibular ligaments (ATFL), 122, 311 312, 420, 510 Anterior tibiotalar ligaments (ATTL), 420 Anteroposterior view (AP view), 690 691 Anthropoid, 701 Applied studies of orthoses and shoe conditions, 173 174 Arch height, 600 Arch height index (AHI), 454 Arch index, 453 Arteries, 41 42 Arthritis, 738 at distal joints, 738 739 Arthrodesis, 100, 702, 721 of calcaneocuboid joint, 714 Articular cartilage, 92, 418, 596 Articulated custom ankle-foot orthoses, 692 As low as reasonably achievable (ALARA), 180 Assessment of foot kinematics, 479 480 Associated ankle pathology, treatment of, 698 ASTM. See American Society for Testing and Materials (ASTM) ATFL. See Anterior talofibular ligaments (ATFL) Athletic shoes, 623 Atrophy, 435 ATTL. See Anterior tibiotalar ligaments (ATTL) Augmented reality systems (AR systems), 236 Autograft wedges, 676 677 Axial plane. See Transverse plane

B Balance, 105, 534 535 Barefoot, 618 biomechanics of, 626 628 running, 616, 618 comparison of full minimal to, 629 comparison of mechanics between, 626 628 pattern, 626 Beams, 727 728 Beuchel-Pappas TAR (BP TAR), 734 735 Biological tissues future biomechanics research, 426 limitations of computational modeling, 425 426 materials and methods, 412 426 numerical analyses of foot functionality, 421 425 BiOM T2, 755 Biomechanical testing, 727 Biomechanics, 45, 485 of athletic footwear, 611 anatomy of running shoe, 612 design, 612 615 footwear related injuries, 619 future footwear research, 620 graphene outsoles, 618 shod vs. barefoot, 618

shoe innovations, 617 618 types of shoes and features, 615 618 dynamics, 48 metrics, 757 of pathologic foot, 373 statics, 46 strength of materials and deformation, 48 49 terminology, 45 46 viscoelasticity, 49 50 Biometrics, 294 295 Biphasic theory, 94 Biplanar linear EOS system, 274 Biplane fluoroscopy, 84 85, 179, 640 background and history, 180 181 challenges specific to foot and ankle tracking with, 182 future applications and directions, 190 191 techniques for tracking foot bone kinematics, 181 182 Biplane hardware, 182 185 Biplane software, 185 187 Biplane systems, 179 C-arms, 187 188 history and evolution, 180 181 independent X-ray sources and image intensifiers, 189 190 modified C-arms, 189 Blood supply, 41 44 arteries, 41 42 lymphatics, 43 44 veins, 42 Body mass index (BMI), 327, 433, 552, 731 732 Bone mineral density (BMD), 91, 595 Bone volume fraction (BV/TV), 91 Bone(s), 305 307, 380, 573, 595 components and structure, 89 92 contusions, 307 density properties, 283 fracture, 373 groupings, 152 153 and other internal anatomy, 344 345 screws, 722 Bony anatomy with arthroplasty, 737 forefoot external rotation, 674 675 hindfoot valgus, 670 671 medial arch eversion, 682 sag at talonavicular joint, 675 Bony defects and ligament injuries, 313 Bony morphology and alignment with arthrodesis, 738 Bony talonavicular external rotation, 675 Botulinum toxin A (BTX-A), 494 injection, 540 BP TAR. See Beuchel-Pappas TAR (BP TAR) Bracing, 405, 692 Bruening kinetic model, 157 Bumpers, 752 Bunions, 527 Bursa/synovia, 307 308

Index

C CAD software. See Computer aided design software (CAD software) Cadaveric flatfoot deformity model, 675 676 Cadaveric gait simulation, 351. See also Dynamic gait simulation areas of future biomechanical research, 360 of arthroplasty and arthrodesis, 739 740 clinical applications of dynamic gait simulation, 359 360 limitations of dynamic gait simulation, 354 359 techniques for dynamic gait simulation, 352 354 Cadaveric gait simulators, 248, 359 360 Cadaveric simulators, 248 249 CADENCE, 735 CAI. See Chronic ankle instability (CAI) Calcaneal fractures, 464 465, 706. See also Pilon fractures diagnostics/classification, 464 etiology and pathophysiology, 464 465 symptoms, 464 treatment, 464 465 goals in, 712 Calcaneal osteotomy, 695 Calcaneal pitch angle, 708 Calcaneal tuberosity, 671 Calcaneocuboid distraction arthrodesis, 675 676 Calcaneocuboid fusions, 714 715 Calcaneocuboid joint, 14 17, 79, 98, 701, 717 Calcaneofibular ligament (CFL), 122, 311, 420, 510 Calcaneus, 6 7, 464, 716 Callus patterns, 436 437 Calluses, 689 Cancellous bone, 89 90, 305 Cannulated screws, 723 724 Cardinal planes, 62 Cartilage, 92 95, 380, 573, 596 ECM, 92 Cast shoes, 662 Casting method, 130 131 Casual shoes, 615 Cavo-varus foot. See Pes cavus, foot Cavovarus deformity, 705 Cavovarus feet, 707 Cavus alignment, 687 688, 690 691 Cavus deformities, 688 Cavus feet, 451 Cavus foot alignment, 692 clinical presentation and associated pathology, 688 689 conservative management, 692 deformity, 687 688 etiology, 687 688 idiopathic, 688 neurologic, 687 688, 688t residual clubfoot, 688 traumatic, 687 imaging, 690 691

biomechanical changes of pes cavus, 691 692 weight-bearing CT scan, 691 X-rays, 690 691 physical exam, 689 surgical management, 693 699 calcaneal osteotomy, 695 complications, 699 dorsiflexion osteotomy or fusion of first ray, 694 gastrocnemius recession/Achilles tendon lengthening, 693 694 lesser toe deformities, 696 modified Jones procedure, 696 overview, 693 peroneus longus to brevis transfer, 694 plantar fascia release, 694 posterior tibial tendon lengthening and transfer, 694 postoperative care, 698 treatment of associated ankle pathology, 698 triple arthrodesis and midfoot fusion, 697 698 syndromes, 705 type, 439 CBCT. See Cone beam CT (CBCT) CDOs. See Custom dynamic orthoses (CDOs) Center of pressure (CoP), 247, 582, 602 603, 755 756 Center of pressure excursion index (CPEI), 247, 454 Centers for Medicare and Medicaid Services, 753 Central activation ratio (CAR), 217 Cerebral palsy (CP), 81, 492 494, 651 structural deformities and gait deviations, 493 treatment, 494 Cervical ligament (CL), 124 CFL. See Calcaneofibular ligament (CFL) Charcot foot. See Charcot neuroarthropathy Charcot neuroarthropathy, 499 500, 568 570, 727 Charcot-Marie-Tooth disease (CMT disease), 483, 494, 500 502, 687 688, 701 702 Charged-coupled devices (CCDs), 265 266 Chemical fat suppression, 302 Children, 477 478 flatfoot, 206 Chondrocytes, 92 93 Chopart’s joint. See Transverse tarsal joint Chronic ankle instability (CAI), 105, 402, 509 513 anatomy, 510 clinical impairments, 512 balance, 512 functional activity, 512 ROM, 512 strength, 512 treatment, 512 513 etiology, 510 512

767

mechanism of injury and pathomechanics, 510 512 Chronic injuries, 507 areas of future research for chronic foot and ankle injuries, 520 523 CAI, 510 513 impairment-based rehabilitation model for treating chronic injuries, 509 through macrotrauma, 509 through microtrauma, 507 509 patient-oriented outcomes, 510 plantar fasciitis, 513 514 retrocalcaneal bursitis, 520 sesamoiditis, 518 520 stress fractures, 516 518 tendinopathy, 514 516 Chronic stress aberration, 399 402 Circulation, 444 Classic chemical fat saturation, 302 Claw toe deformities, 691 Clinical biplane foot and ankle studies, 187 190 Clinical examination of foot and ankle, 433 areas of future biomechanical research, 446 assessment of pain, 434 circulation, 444 demographics, 433 foot and ankle specific testing, 444 foot posture or foot shape, 438 439 footwear examination, 444 functional assessment, 445 joint mobility, 441 442 ligamentous/stability testing, 442 limb length, 440 lower extremity alignment, 438 muscle strength, 443 outcomes assessment, 445 446 patient history, 433 434 radiographic examination, 440 range of motion/flexibility/joint mobility, 440 441 sensory testing, 443 tendon, 442 visual observation/inspection, 434 438 vitals signs, 433 Clinical trials, 756 758 examining effects of damping, 757 examining effects of stiffness, 756 757 Clubfoot, 479 CMT joint. See Cuneiometatarsal joint (CMT joint) Cole midtarsal osteotomy, 698 Coleman block test(ing), 689, 707 Collagen, 598 Collimators, 182 183 Community-based video task analysis, 752 Comparative effectiveness, 743 Complex hindfoot biomechanics, 702 705 Compression, 721 722 plate, 724 Computational modeling, 412 Computational models, 283 Computed muscle control, 390 Computed radiography (CR), 265 266

768

Index

Computed tomography (CT), 84, 91, 179, 186 187, 277, 289, 378, 388, 532 533, 706 areas of future biomechanical research, 284 285 comparison to other imaging modalities, 278 279 foot-specific applications and considerations, 281 284 biomechanics research, 282 284 disease diagnosis, 281 surgical assessment and planning, 281 282 history and development, 277 278 images, 412 imaging, 452 MRI vs., 303 304 protocols for foot and ankle, 279 280 scans, 247, 461 462, 708 Computer aided design software (CAD software), 341 Cone beam CT (CBCT), 278, 290, 293 Congenital foot deformities, 479 482 Connective tissue disorders, 591 Conservative treatment, 534 Constitutive formulation, 412 Constitutive models, 412 formulation, 412 416 identification of constitutive parameters, 417 420 ankle ligaments, 420 bone, 417 cartilage, 418 plantar soft tissue, 419 420 Contact scanners, 341 Conventional AFOs, 655 Conventional double-upright designs, 652 Conventional footwear, comparison of full minimal to, 630 Conventional shod comparison of mechanics between, 626 628 running, 626 628 pattern, 626 Coordinate systems, 153 CoP. See Center of pressure (CoP) Coronal plane, 62 Cortez shoe, 624 Cortical bone, 89 90, 305, 417, 595 “Cortical” screws, 723 Cotton osteotomy, 682 CP. See Cerebral palsy (CP) CPEI. See Center of pressure excursion index (CPEI) Cross-sectional area (CSA), 104 Cuboid, 7 Cuboideonavicular joint, 17 Cueing, 237 Cuneiometatarsal joint (CMT joint), 552 Cuneocuboid joints, 17 18 Cuneonavicular joints, 17 Cushioned shoes, 624 Cushioning, 207, 613 614 Custom dedicated biplane hardware for foot bone tracking, 181

Custom dynamic orthoses (CDOs), 405 Custom-made insoles, 662 663

D Damping, examining effects of, 757 Darkroom, 265 266 Deep posterior compartment syndrome, 687 Deep veins, 42 Deformation, 48 49 Degenerative joint disease, 315 316 Degenerative phenomena, 411 Degree of anisotropy, 91 Depth cameras, 342 Development of normal foot progression angle, 477 Developmental foot deformities, 482 DEXA. See Dual energy X-ray absorptiometry (DEXA) DGSs. See Dynamic gait simulators (DGSs) Diabetes, 205 206 Diabetic condition, 425 Diabetic foot disease, 371, 565, 661 areas of future biomechanical research, 575 changes in kinematics and kinetics in, 571 573 tissue characteristics, 573 574 key negative outcomes, 566 571 relationship between foot deformities and plantar ulceration, 574 relationship between lower extremity fractures and Charcot neuropathic osteoarthropathy, 574 risk factors for development and progression, 571 Diabetic footwear biomechanical effect, 662 663 foot biomechanics and offloading, 661 662 footwear and offloading adherence, 665 666 footwear and offloading for ulcer healing, 663 664 for ulcer prevention, 664 665 Diabetic neuropathy, 113 Diabetic peripheral neuropathy (DPN). See Distal sensory polyneuropathy (DSP) Diabetic plantar soft tissue, 144 other pathologies associated with plantar soft tissue, 144 Diabetic plantar ulceration, 567 Diabetic polyneuropathy (DPN), 113 Digital radiography (DR), 265 266 Digitally reconstructed radiographs (DRRs), 185 186, 189, 292 Dijian-Annonier angle, 270 Diplegia, 651 Disarticulated C-arm systems for foot bone tracking, 181 Distal intermetatarsal joints, 19 Distal joints, arthritis at, 738 739 Distal phalanges, 11 Distal sensory polyneuropathy (DSP), 495 496, 566 567

Distraction subtalar fusion, 712 713 Dorsal interossei, 37 38 Dorsiflexion, 422, 491 492 osteotomy of first ray, 694 Double hindfoot arthrodesis, 714 715 Double hindfoot fusion, 714 715 Double hump pattern, 203 Doublet, 216 Dual energy X-ray absorptiometry (DEXA), 91, 283 Duchenne musculardystrophy (DMD), 494 Dynamic gait simulation clinical applications of, 359 360 limitations of, 354 359 techniques for, 352 354 Dynamic gait simulators (DGSs), 351 352 Dynamic(s), 48 3D scanning, 342 arch index, 454 navicular drop, 454 paradigm, 218 219 tendon force, 352 Dynamometry, 214 219 dynamic paradigm, 218 219 isometric myograph, 214 218

E Early invasive studies of foot and ankle biomechanics, 168 169 Early rheumatoid arthritis, 582 Edema, 435, 708 Elastic constants, 415 Elastic deformation, 49 Elastic mechanisms, 702 Elastic modulus, 49 Elastography, 324 Elastomeric springs, 752 Electric strain gages, 232 Electromagnetic energy, 265 266 Electromyography (EMG), 211 214, 389 390, 529, 642 activity, 104 indwelling electromyography and motor unit recordings, 213 214 surface electromyography, 212 213 Empower device, 755 End-stage ankle arthritis (ESAA), 731 732 Energy absorption, 140 Energy storing and returning (ESAR), 652 653, 752 prosthetic feet, 760 Enthesitis, 315 Epidemiology, 246 case study of integrated laboratories concept to study of hallux rigidus, 253 integrated laboratories, 247 Eponyms, persistence of, 54 Equino-varus foot, 651 Equinus contracture, 710 Essential joints, 701 Established rheumatoid arthritis, 582 588 Evans LCL, 676 677 Excessive hindfoot valgus, 718

Index

Exercise, 540, 549 Exostoses, 437 Expansion joint, 703 Experimental testing, 412 Explicit solvers, 382 Extensor digitorum brevis (EDB), 29 30, 103 Extensor digitorum longus (EDL), 22 24, 103 Extensor hallucis brevis (EHB), 29 30 Extensor hallucis longus (EHL), 24 25, 103, 694 External fixation, 733 External-internal rotation, 67 “Extra medullary” plates, 727 Extracellular matrix (ECM), 89 Extrinsic dorsal muscles, 21 25 extensor digitorum longus, 22 24 extensor hallucis longus, 24 25 peroneus tertius, 25 tibialis anterior, 21 Extrinsic foot muscles, 103, 630 Extrinsic lateral muscles, 28 29 peroneus brevis, 29 peroneus longus, 28 29 Extrinsic plantar muscles, 25 28. See also Intrinsic plantar muscles flexor digitorum longus, 26 flexor hallucis longus, 26 27 plantaris, 26 tibialis posterior, 27 28 triceps surae, 25 26

F Failure, 49 Falls, 534 535, 604 FAO. See Foot ankle offset (FAO) FAP. See Foot anatomical position (FAP) Fascial specializations, 21 Fasciitis, 326 327 Fast spin echo (FSE), 302 FastSCAN, 342 Fat pad, 599 Fatty infiltration and reduction of intrinsic foot muscle volumes, 573 FCI. See Functional Comorbidity Index (FCI) FDB. See Flexor digitorium brevis (FDB) FDL. See Flexor digitorum longus (FDL) FE modeling. See Finite element modeling (FE modeling) Feet, 451 FFS. See Forefoot strike (FFS) FHB. See Flexor hallicus brevis (FHB) FHL. See Flexor hallucis longus (FHL) Fibrocartilaginous cartilage, 307 Fibula, 3 Fibular nerves in foot, 38 39 Field of view (FOV), 279, 291 292 Fine-wire technique, 213 Finite element modeling (FE modeling), 249, 251, 365, 390, 740 anatomically detailed compared to idealized modeling, 378 379 applications of finite element analysis in foot biomechanics, 366 374

simulation of healthy foot biomechanics, 369 371 simulation of interaction between foot and footwear, 366 369 basic concepts, 365 366 of foot, 412 for in vivo material characterization of soft tissues, 371 372 limitations and future research, 383 modeling of entire foot compared to anatomically focused modeling, 377 modeling strategies, 375 382 geometry design, 375 379 material properties, 380 381 meshing, 379 380 reliability assessment, 382 solver selection, 381 382 simulation of pathological conditions, 373 374 First metatarsophalangeal joint osteoarthritis, 549 552 clinical and biomechanical effects of conservative treatment, 551 552 clinical findings, 549 550 etiology and impact, 549 structural and biomechanical features, 550 551 Fitness shoe, 654 655 Fixed-angle stiffness prosthetic foot, 753 754 Flat panel detectors (FPD), 182 Flatfoot, 705 reconstruction, 672, 682 treatment goals in flatfoot/cavus syndromes, 710 Flexible claw toes, 696 Flexible flatfoot, 482 deformity, 705 Flexion-extension, 65 66 Flexor accessories. See Quadratus plantae Flexor digiti minimi brevis, 36 Flexor digitorium brevis (FDB), 33, 103 Flexor digitorum longus (FDL), 26, 313 314, 669 670 tendon transfer, 680 Flexor hallicus brevis (FHB), 35 36, 103 Flexor hallucis longus (FHL), 26 27, 103, 313 314 Foot, 77 78, 289, 477, 643 biomechanics, 425 426, 661 662 bone kinematics, 181 182 comorbidity, 126 deformity, 482 483, 527, 568, 603 disorders, 603 604 drop, 496 498, 651 foot-specific applications and considerations, 270, 342 345 bones and other internal anatomy, 344 345 musculoskeletal models, 344 orthoses and footwear, 344 population studies, 343 344 reliability and comparisons to other techniques, 342 and footwear interaction, 424

769

function, 601 603 kinematics, 603 kinetics, 601 603 midline, 64 effect of minimal shoes on foot musculoskeletal system, 630 631 morphology, 344 motions, 64 71 whole foot motions and complexity, 67 71 offloading, 661 662 pain, 534 pathologies associated with other conditions, 482 484 posture, 603 604 anthropometrics, 599 arch height, 600 cavus foot type, 439 joint range-of-motion, 600 601 and morphology, 599 601 or foot shape, 438 439 planus foot type, 439 stability, 104 structure measures, 247 symptoms, 603 604 type biomechanics, 454 456 areas of future biomechanical research, 457 association with pain and injury, 456 457 functional foot type, 454 structural foot type, 452 454 treatments, 457 Foot anatomical position (FAP), 56 Foot and ankle basic research in walking and running kinematics, 173 biomechanics of foot and ankle fixation beams, 727 728 nails, 726 727 plates, 724 726 post and screw constructs, 726 screws, 722 724 osteoarthritis subtypes, 549 specific testing, 444 Foot ankle offset (FAO), 294 295 Foot Function Index scores, 651 652 Foot orthoses (FO), 539, 637 biomechanical effects of, 639 642 effects on muscle activity patterns, 642 effects on plantar pressure, 641 kinematic effects of, 640 641 kinetic effects of, 641 design and manufacture of, 637 639 effects on clinical conditions, 642 643 heel pain, 643 osteoarthritis, 643 rheumatoid arthritis, 642 643 sports injuries and conditions, 643 symptomatic flat foot, 643 Foot posture index (FPI), 453 Footwear, 113 114, 344, 366 369, 534, 611 adherence, 665 666 companies, 624

770

Index

Footwear (Continued) embedded energy harvester, 617 examination, 444 interventions, 647 modalities, 664 related injuries, 619 for ulcer healing, 663 664 Force, 45 plates, 203 platforms, 199, 203 sensing, 231 232, 234 235 Forefoot, 55, 126, 317 318, 692 external rotation, 674 677 bony anatomy, 674 675 ligament and tendon failure, 675 surgical reconstruction, 675 677 forefoot-driven cavus, 692 Morton’s neuroma, 317 offloading shoes, 662 osteomyelitis, 318 plantar plate tears, 318 rocker, 649 Forefoot driven hindfoot varus, 692 Forefoot strike (FFS), 107, 626 Forward dynamics, 390 Four-dimensional CT (4DCT), 284 285 FOV. See Field of view (FOV) FPD. See Flat panel detectors (FPD) Fracture, 461 repair, 721 severity, 398 Frataxin (FXN), 502 Free-body diagrams, 46 Friederichs’s ataxia (FRDA), 502 503 Frontal plane. See Coronal plane FSE. See Fast spin echo (FSE) Full minimal footwear, 628 629 Full minimal to barefoot running, comparison of, 629 Full minimal to conventional footwear, comparison of, 630 Full minimal to partial minimal shoes, comparison of, 629 630 Full minimalist footwear, definition of, 628 629 Fully-threaded nonlag screws, 713 Functional Comorbidity Index (FCI), 743 Functional foot type, 454 Fusion of first ray, 694

G Gadolinium contrast media, 301 302 Gait, 105, 479 480 analysis, 247, 535 536 kinematics, 535 536 muscle activity, 536 plantar pressures, 536 temporospatial parameters, 536 cycle, 77 78, 423 424 deficiency, 651 deviations, 651 652 mechanics, 738 Gastrocnemius equinus, 692

Gastrocnemius recession/Achilles tendon lengthening, 693 694 Gastrocnemius tightness, 680 681 Gauntlet, 652 Gel pockets, 639 Geometric measurements, 270 Geometry reconstruction, 377 Glasgow Maastricht foot model, 392 Glycosaminoglycans (GAGs), 92 Gout, 591 592 Graphene outsoles, 618 Ground reaction forces (GRFs), 80, 249, 454, 601, 613 altering line of action of, 648

H HA. See Hydroxyapatite (HA) Habitual motion path, 620 Hallucal metatarsophalangeal joint, 19 20 Hallux flexion and abduction, 535 Hallux rigidus (HR), 271 case study of integrated laboratories concept to study of, 253 260 epidemiology, 253 in silico simulation, 258 260 in vitro experimentation, 256 258 in vivo experimentation, 254 256 Hallux valgus (HV), 64, 527 clinical presentation, 534 diagnosis and imaging, 529 534 clinical diagnosis, 529 530 computed tomography, 533 magnetic resonance imaging, 533 534 radiographic assessment, 530 532 ultrasound, 532 533 etiology, 528 529 footwear, 529 genetics and race, 528 structural and biomechanical factors, 528 529 functional outcomes, 534 536 future directions for research, 541 prevalence, 527 treatment pathways, 536 540 botulinum toxin A injection, 540 exercise, 540 expert opinion and current practice, 537 538 foot orthoses, 539 manual therapy, 540 nonsurgical treatment, 536 540 splints and toe separators, 539 surgical treatment, 540 taping, 540 Hallux valgus angle (HVA), 270 Hammer toe deformities, 691 HAP. See Human anatomical position (HAP) Haptics, 236 237 cueing and notification, 237 sensory substitution, 236 237 HCPCS. See Medicare Healthcare Common Procedure Codes System (HCPCS) Healing plantar foot ulcers, 663

Health and activity monitoring, 235 sensing, 232 233 Heel pad, 139 Heel pain, 371, 643 Heel rocker, 649 Heel strike forces, 741 Helmholtz free energy function, 416 Hemiplegia, 647 Hemiplegic gait, 650 651 Hereditary motor and sensory neuropathies (HMSN), 687 688 Hibb’s angle, 270 High ankle sprains. See Tibiofibular syndesmosis, sprains High arch foot mechanics, 687 Hindfoot, 55, 313 315, 691 692 alignment, 707 arthrodesis, 701 702 driven cavus, 691 692 driven deformities, 691 692 enthesitis, 315 malalignment, 710 stability, 614 615 tendons, 313 315 varus, 688 689 Hindfoot fusions complex hindfoot biomechanics, 702 705 conditions require, 705 707 accessory navicular, 707 calcaneal fractures, 706 cavus foot syndromes, 705 flatfoot, 705 osteoarthritis, 706 rheumatoid arthritis, 706 talar fractures and dislocations, 706 tarsal coalitions, 706 corrective options, 712 718 distraction subtalar fusion, 712 713 double hindfoot fusion, 714 715 general complications, 718 lateral column lengthening, 713 714 pantalar arthrodesis, 717 718 postoperative management requirements in hindfoot fusions, 718 subtalar and talonavicular fusion, 715 716 triple arthrodesis, 716 717 goals in treatment, 710 712 treatment goals in calcaneal fractures, 712 treatment goals in flatfoot/cavus syndromes, 710 treatment goals in osteoarthritis, 712 treatment goals in rheumatoid arthritis, 711 712 treatment goals in talar fractures and dislocations, 712 imaging, 708 710 presurgical assessment, 707 clinical exam, 707 Hindfoot ligaments, 124. See also Midfoot ligaments subtalar joint, 124 talonavicular joint, 124

Index

Hindfoot valgus, 670 673, 677 678 bony anatomy, 670 671 deformity, 671 ligament failure, 671 surgical reconstruction, 672 673 Histology, 248 History-dependence of force, 221 224 HMSN. See Hereditary motor and sensory neuropathies (HMSN) Horizontal plane. See Transverse plane HR. See Hallux rigidus (HR) Human anatomical position (HAP), 55 56 application specific, 56 61 HV. See Hallux valgus (HV) HVA. See Hallux valgus angle (HVA) Hyaline cartilage, 307 Hydroxyapatite (HA), 89 Hyperelastic constitutive models, 415 416 Hysteresis, 200

I Idiopathic toe walking (ITW), 494 Image segmentation, 250 Impact forces, 613 614 Impairment-based rehabilitation model for treating chronic injuries, 509 Implicit solvers, 381 In silico simulation, 246, 249 253 finite element modeling, 251 medical image processing, 250 251 musculoskeletal modeling, 252 sensitivity studies, 252 validation, 252 253 In situ fusions, 710 In vitro experimentation, 246, 248 249 cadaveric simulators, 248 249 histology, 248 In vivo experimentation, 246 247 foot structure measures, 247 gait analysis, 247 other measures, 247 plantar pressure assessments, 247 In vivo techniques, 143 In-shoe plantar pressure analysis, 664 665 In-shoe pressure measurement systems, 641 In-shoe systems, 198 INBONE II, 735 Incongruency angle, 674 675 Indwelling electromyography, 213 214 Inertial measurement unit (IMU), 235 Inferior calcaneonavicular component, 678 INFINITY, 735 Insertion of markers in bones of foot and ankle, 175 176 Insertional Achilles tendinopathy (IAT), 330 Inserts and bracing, 692 Insoles, 637 Intact C-arm systems for foot bone tracking, 181 Integrated laboratories, 247 253 case study of integrated laboratories concept to study of hallux rigidus, 253 260

epidemiology, 247 in vitro experimentation, 248 249 in vivo experimentation, 247 Intercuneiform joints, 17 18 Interlocking screws, 727 Intermediate cuneiform, 7 8 Intermetatarsal angle (IMA), 271 Internal fixation, 721 of foot fractures, 724 plates, 724 techniques, 721 Internal fixators, 726 Internal reinforcing beams, 728 International Organization for Standardization (ISO), 758 International Society of Biomechanics (ISB), 56 61 International Working Group on Diabetic Foot (IWGDF), 665 Interphalangeal joints (IP joints), 20, 691 Interpolated twitch technique (ITT), 217 Intracalcaneal lengthening osteotomy, 711 712 Intraclass correlation coefficients (ICC), 255 256 Intracortical pins, 167 applications and significance of studies for foot and ankle kinematics, 171 174 ankle kinematics, 173 applied studies of orthoses and shoe conditions, 173 174 foot and ankle basic research in walking and running kinematics, 173 skin movement artifact in foot and ankle kinematics, 171 172 Intramedullary nails, 726 727 Intramuscular EMG measurements, 390 techniques, 213 Intrepid Dynamic Exoskeletal Orthosis device design, 405 Intrinsic dorsal foot muscles, 29 30 extensor hallucis brevis and extensor digitorum brevis, 29 30 Intrinsic foot and ankle kinematics, 167 muscles, 103 Intrinsic forces, 388 Intrinsic plantar muscles, 31 38. See also Extrinsic plantar muscles abductor digiti minimi, 33 abductor hallucis, 33 adductor hallucis, 36 37 dorsal and plantar interossei, 37 38 flexor digiti minimi brevis, 36 flexor digitorum brevis, 33 flexor hallucis brevis, 35 36 lumbricals, 34 35 quadratus plantae, 33 Invasive methods for ankle and intrinsic foot kinematics, 167 applications and significance of studies using intracortical pins for foot and ankle kinematics, 171 174

771

early invasive studies of foot and ankle biomechanics, 168 169 limitations and future directions, 174 175 radiostereometric analysis, 169 171 Inverse dynamics, 389 390 Inverse engineering, 376 Inversion eversion, 68 69 Involuntary isometric contractions, 215 Isokinetic contractions, 219 Isolated fusion of subtalar joint, 697 Isolated subtalar fusion, 712 Isolated talonavicular arthrodesis, 711 712 Isometric MVC, 216 217 Isometric myograph, 214 218 Isotonic contractions, 219 Isotropic models, 414 415 36-Item Short Form survey (SF-36), 510 ITW. See Idiopathic toe walking (ITW)

J Jahss truncated wedge osteotomy, 698 Japas midtarsal osteotomy, 698 Joint(s), 11 20 functions, 95 100 injury, 397 mobility, 441 442 moments and powers, 601 range-of-motion, 600 601 space narrowing, 682 tibiofibular syndesmosis, 11 Juvenile bunion, 482 Juvenile idiopathic arthritis (JIA), 581, 590 591

K K-levels, 750 Kellgren-Lawrence scores (KL scores), 398 399 Kinematics, 48, 79 80, 151, 351 352, 389, 479 480, 552 future areas of research, 84 85 healthy and impaired feet, 81 measurements, 282 multisegment foot models, 81 84 Kinesiology, 48 Kinetics, 48, 151, 351 352, 389 390, 557 modeling, 79 80 future areas of research, 84 85 healthy and impaired feet, 81 multisegment foot models, 81 84 Knee, 643 angle, 221

L Lacing systems, 617 618 Lag screw, 726 Lamellae, 417 Lateral collateral ligaments (LCL), 122 Lateral column lengthening (LCL), 669 670, 713 714 Lateral cuneiform, 8

772

Index

Lateral projection, 708 Lateral radiographs, 682, 708 Lateral subluxation of calcaneus, 671 Lateral talometatarsal angle, 708 Lateral weightbearing radiographs, 677 678 Lateralizing calcaneal osteotomy, 695 Leardini/Rizzoli kinematic model, 155 156 Leg Anatomical Position (LAP), 56 61 Lesser metatarsophalangeal joints, 19 Lesser toe deformities, 696 “Ligament-compatible” ankle replacement, 737 Ligamentous ankle anatomy with arthroplasty, 737 Ligamentous tissue, 420 Ligamentous/stability testing, 442 Ligaments, 121, 305, 311 313, 380 381, 598 anatomy, 121 122 future areas of research, 132 mechanical properties, 122 127 overcoming limitations, 127 131 sprains, 127 variations in mechanical properties, 126 127 age effects, 127 changes in activity level, 126 foot comorbidity, 126 influence of anthropometric effects, 127 Limb alignment with arthroplasty, 736 737 Limb length, 440 Limb loss, 749 Limb salvage, 652 Linear elastic constitutive models, 414 415 Linearity, 200 Lisfranc’s joint. See Tarsometatarsal joints Load distribution, 140 Locking screws, 725 Locomotion, 105 108 Loss of protective sensation (LOPS), 661 Low-grade lateral process, 708 Lower extremity alignment, 438 amputation, 570 571 Lower limb distal element of, 54 55 loss, 750 prosthesis, 750 Lower-extremity fractures, 568 Lumbricals, 34 35 Lymphatics, 43 44

M Macrotrauma, chronic injury through, 509 MacWilliams/Kinfoot kinetic model, 157 Magic angle effect, 305 Magnetic resonance (MR), 412 Magnetic resonance imaging (MRI), 104, 140, 143, 179, 249, 278 279, 301, 367, 388, 529, 533 534, 584, 706 anatomy of foot and ankle, 311 318 appearance of musculoskeletal tissue, 304 308 areas of future research, 318 320 computed tomography vs., 303 304

forefoot, 317 318 hindfoot, 313 315 midfoot, 315 317 scan time reduction, 319 sequences, 302 303 tailored magnetic resonance imaging protocol for foot and ankle, 308 311 imaging orientation, 308 metal artifact reduction sequences, 311 optimized imaging planes, 310 311 tailored magnetic resonance imaging protocols, 308 309 Malerba Z-cut osteotomy, 695 Manchester Scale, 529 Manual therapy, 540 Marathon shoes, 616 Marker type and placement, 153 Marker-based tracking software, 185 186 Masai BarefootTechnology (MBT), 654 655 Massachusetts General Hospital (MGH), 187 Mat systems, 198 Material ex vivo testing, 142 MAVRIC. See Multiacquisition variableresonance imaging combination (MAVRIC) Maximal voluntary contractions, 221 Maximal voluntary isometric contraction (MVC), 213 Meary’s angle, 677 678 Mechanical property tests, 758 759 roll-over shape, 759 stiffness, hysteresis, and energy, 758 759 Mechanical testing, 127 Mechanoreceptors, 598 Medial arch eversion, 681 682 Medial collateral ligament (MCL), 123 Medial cuneiform, 7 Medial displacement calcaneal osteotomy, 711 712 Medial hindfoot posting, 640 641 Medial keratotic lesions, 705 Medial longitudinal arch (MLA), 452f, 681, 703 Medializing calcaneal osteotomy (MCO), 669 670 Medical image processing, 250 251 Medical imaging, 279 Medicare functional classification levels (MFCL), 750 Medicare Healthcare Common Procedure Codes System (HCPCS), 752 753 Meshing, 379 380 element type selection, 380 mesh convergence, 380 Metabolic energy expenditure effects, 757 Metal artifact reduction algorithms (MAR algorithms), 279 280, 311 Metatarsals, 8 10, 90 base ligaments, 124 125 fifth, 9 10 first, 8 9 fourth, 9 fractures, 468 472 second, 9

third, 9 Metatarsocuneiform (MTC), 253 Metatarsophalangeal joints (MTP joints), 79, 99 100, 245, 491 492, 513, 527, 549, 616, 689 Metatarsus, 8 Metatarsus primus elevatus (MPE), 271 Micro-CT scanners, 91 Micromotion, 721 Microscopy, 230 Microtrauma, chronic injury through, 507 509 Middle phalanges, 10 Midfoot, 315 317, 468 arthrodesis, 679 crush injuries, 472 degenerative joint disease, 315 316 fusion, 697 698 locking mechanism, 98 99 stress fractures, 316 317 Midfoot ligaments, 124 125. See also Hindfoot ligaments metatarsal base ligaments, 124 125 plantar fascia, 124 Midfoot osteoarthritis, 552 555. See also Ankle osteoarthritis clinical and biomechanical effects of conservative treatment, 555 clinical findings, 553 etiology and impact, 552 structural and biomechanical features, 553 555 Mild-to-moderate deformities, 693 Milwaukee Foot Model (MFM), 84 Milwaukee kinematic model, 154 Minimal footwear running, 628 630 comparison of full minimal to barefoot running, 629 comparison of full minimal to conventional footwear, 630 comparison of full minimal to partial minimal shoes, 629 630 definition of full minimalist footwear, 628 629 Minimal shoes biomechanics of barefoot and conventional shod running, 626 628 barefoot running pattern, 626 comparison of mechanics between conventional shod and barefoot running, 626 628 conventional shod running pattern, 626 brief history of running footwear, 623 625 minimal footwear running, 628 630 effect of minimal shoes on foot musculoskeletal system, 630 631 Minimalism, 628 629 Minimalist index, 628 629 MLA. See Medial longitudinal arch (MLA) Mobility limitations, 603 604 Model-based tracking software, 185 186 Modeling platforms, 390 Modified Jones procedure, 696 Molding approach, 130 131, 638 Morton’s neuroma, 317, 499

Index

Motion capture, 151 152, 237 238 Motion-sacrificing procedures, 710 Motor unit (MU), 211 behavior and quantity, 219 220 recordings, 213 214 Motor unit number estimations (MUNEs), 220 MPE. See Metatarsus primus elevatus (MPE) MPJ. See Metatarsophalangeal joints (MTP joints) MRI. See Magnetic resonance imaging (MRI) MTC. See Metatarsocuneiform (MTC) MTP joints. See Metatarsophalangeal joints (MTP joints) Multi planar reconstruction (MPR), 292 Multiacquisition variable-resonance imaging combination (MAVRIC), 311 Multisegment foot models (MFM), 81 84, 152 areas of future biomechanical research, 161 basic principles, 151 152 clinical applications, 159 160 direct comparison of current, 157 159 review, 154 159 selecting, 152 154 sources of error, 160 161 standardized description, 154 Muscle(s), 20 38, 103, 307, 596 597 aging, 108 109 areas of future biomechanical research, 116 117 atrophy, 435 biomechanical function, 104 116 effects of foot orthosis on muscle activity patterns, 642 extrinsic dorsal muscles, 21 25 extrinsic lateral muscles, 28 29 extrinsic plantar muscles, 25 28 footwear and orthoses, 113 114 interventions, 114 116 intrinsic dorsal foot muscles, 29 30 intrinsic plantar muscles, 31 38 pathologies, 109 113 strength, 443 tendons, 3 weakness, 489 Musculoskeletal (MSK), 245 modeling, 387, 756, 759 areas of future biomechanical research, 394 challenges in modeling foot and ankle, 391 392 and development, 388 390 foot specific models and applications, 392 394 validation techniques, 390 models, 344 simulation, 759 system, 245, 249, 252 tissue, 304 308 Musculoskeletal Functional Assessment, 743 MVC. See Maximal voluntary isometric contraction (MVC) Myograph, 214 215

N Nails, 726 727 Navicular, 7 drop, 453 height, 453 Naviculocuneiform joint (NC joint), 552 Nerves, 38 41 fibular nerves in foot, 38 39 tibial nerves in foot, 40 41 Neural, 598 Neurological foot pathology, 489 cerebral palsy (CP), 492 494 Charcot foot, 499 500 Charcot-Marie-Tooth disease, 500 502 foot drop, 496 498 Friederichs’s ataxia, 502 503 Morton’s neuroma, 499 peripheral neuropathy, 495 496 poliomyelitis, 503 504 stroke, 489 492 tarsal tunnel syndrome, 498 toe walking, 494 495 Neuromuscular system, 211 ankle and foot related considerations and insights, 219 224 dynamometry, 214 219 electromyography, 211 214 future research, 224 Neuropathic plantar forefoot ulcers, 666 Neuropathy, 498 Neutralization plates, 724 Noncontact scanners, 341 342 Nonessential joints, 701 Nonessential stability joints, 701 Nonparalytic pes cavus, 710 Normal foot, 104 108 balance, 105 foot stability, 104 locomotion, 105 108 Notification, 237 Numerical analyses of foot functionality, 421 425 ankle movements, 422 423 diabetic condition, 425 foot and footwear interaction, 424 gait cycle, 423 424

O Oblique and vertical talus, 480 Off-the-shelf carbon fiber designs, 652 devices, 637 Offloading, 661 adherence, 665 666 footwear and, 665 666 braces, 648 devices, 662 biomechanical effect of, 662 663 for ulcer healing, 663 664 Offsets, 153 Open reduction and internal fixation (ORIF), 462

773

ORIF. See Open reduction and internal fixation (ORIF) Original equipment manufacturer (OEM), 180 181 Orthema CAD/CAM system, 341 Orthopedics, 289 Orthosis, 113 114, 344, 367, 551 Orthotics, 637 interventions, 647 Orthotropic materials, 414 415 Osteoarthritis (OA) ankle osteoarthritis, 556 558 areas of future biomechanical research, 558 559 first metatarsophalangeal joint osteoarthritis, 549 552 foot and ankle osteoarthritis subtypes, 549 midfoot osteoarthritis, 552 555 risk factors and classification, 549 structural changes, 548 symptoms and diagnosis, 547 treatment goals in, 712 Osteoarthritis (OA), 94, 397, 547, 643, 706, 731 732 Osteochondral lesion (OCL), 403 Osteomyelitis, 318 Osteons, 417 Osteophytes, 682 Osteotomy, 721 Overuse, 514 515 Oxford kinematic model, 156 157

P Pain, 434 and injury, 456 457 Pantalar arthrodesis, 717 718 Paralytic pes cavus, 710 Partial minimal shoes, comparison of full minimal to, 629 630 Partial weight bearing (PWB), 255 256 Pathomechanics, 397 Patient-oriented outcomes, 510 Patient-reported outcome measures (PROMs), 510 Patient-Reported Outcomes Measurement Information System (PROMIS), 445 Patient-tailored approach, 638 639 PD weighting. See Proton density weighting (PD weighting) Peak plantarflexion, 741 Pedal pathologies, 245 future biomechanical research, 260 262 integrated laboratories, 247 253 method of approach, 246 Pediatric foot, 477 areas for future research, 485 486 common pathologies affecting pediatric feet, 478 484 functional assessment of pediatric foot, 484 485 Pedobarography, 197 198 Peripheral neuropathy, 495 496, 568

774

Index

Peripheral neuropathy (Continued) musculoskeletal and movement implications, 496 Peripheral vascular disease, 567 Peroneal tendon, 516 Peroneal view, 310 Peroneus brevis (PB), 29, 103 Peroneus longus (PL), 28 29, 103 to brevis transfer, 694 Peroneus tertius (PT), 25, 103 Pes cavus, 687, 710 biomechanical changes of, 691 692 forefoot, 692 hindfoot, 691 692 soft tissues, 692 foot, 483 Pes planovalgus, 706 Pes planus, 110 111, 710 Phalanges, 10 11 Phantoms, 270 Photogrammetry, 342 Physical casting, 638 Pigmentations, 434 Pilon fractures, 461 464. See also Calcaneal fractures diagnostics/classification, 461 462 etiology and pathophysiology, 461 symptoms, 461 treatment, 462 464 Plantar aponeurosis (PA), 394 Plantar calcaneonavicular ligament. See Spring ligament Plantar fascia, 104 105, 124, 574 release, 694 thickness assessment in relation to weightbearing activities, 326 327 ultrasound elastography to assess mechanical properties, 333 Plantar fasciitis, 109 110, 315, 513 514 anatomical overview, 513 clinical impairments, 514 functional activity, 514 ROM, 514 strength, 514 treatment, 514 etiology, 513 514 mechanism of injury and pathomechanics, 513 514 Plantar fat pad, 139 Plantar interossei, 37 38 Plantar plate tears, 318 Plantar pressure, 478, 567, 601 603 areas of future research, 207 208 assessments, 247 clinical relevance of force and pressure measurements in foot and ankle biomechanics, 197 198 effects of foot orthosis on, 641 force vs. pressure, 198 203 history, 198 199 measurement technologies, 199 201 visualization and analytical options, 201 203 measurement system, 661 662, 666

research applications and selected clinical examples, 204 207 Plantar soft tissue, 135 effect of aging, 143 144 anatomy, 135 140 gross anatomy, 135 136 histological or biochemical, 136 137 medical imaging of tissue thickness, 137 140 areas of future biomechanical research, 145 biomechanical function, 140 constitutive parameter identification for, 419 420 diabetic plantar soft tissue, 144 mechanical properties, 140 143 thickness assessment in relation to weightbearing activities, 325 326 ultrasound combined with load cells to assess mechanical properties of, 328 329 ultrasound elastography to assess mechanical properties of, 331 332 Plantarflexion, 422 Plantarflexion-dorsiflexion, 67 68 Plantaris, 26 Planus feet, 451 452 foot type, 439 Plates, 724 726 Platform systems, 198 Poisson noise process, 183 184 Poisson’s ratio, 49 Poliomyelitis, 503 504 Polyarticular foot OA, 552 Polypropylene, 637 638 Population studies, 343 344 Post-polio syndrome (PSS), 503 Posterior inferior tibiofibular ligament (PITFL), 311 Posterior leaf springs, 652 Posterior talofibular ligaments (PTFL), 122, 311 312, 420, 510 Posterior tibial tendon (PTT), 675, 703 failure of, 679 680 clinical assessment, 680 MRI assessment, 680 surgical reconstruction, 680 lengthening and transfer, 694 Posterior tibial tendon dysfunction (PTTD), 81, 515 516 Posterior tibialis, 313 314 Posterior tibiotalar ligaments (PTTL), 420 Posterior anterior, 61 Posterolateral approach, 712 713 Postoperative care, 698 Posttraumatic ankle osteoarthritis, 404 407 Posttraumatic osteoarthritis (PTOA), 397. See also Rheumatoid arthritis (RA) acute joint injury severity, 397 399 altered kinematics, 402 404 areas of future biomechanical research, 404 407 chronic stress aberration, 399 402 Powered prosthesis, 757 758

Powered prosthetic feet, 755, 757 758 Preferred movement path, 620 Preferred walking speed effects, 757 Pressure measurement systems, 198 Pressure sensing, 231 232, 234 235 Presurgical assessment, 707 Primary bone healing, 721 PROMIS. See Patient-reported outcomes measurement information system (PROMIS) PROMs. See Patient-reported outcome measures (PROMs) Pronation supination, 70 71 Prosthetic ankle-foot roll-over shapes, 759 Prosthetic feet, 749 form of, 752 755 fixed-angle stiffness, 753 754 powered prosthetic feet, 755 solid ankle cushioned heel, 753 variable stiffness, 755 variable-angle stiffness, 754 755 function of, 755 759 clinical trials, 756 758 future prosthetic foot research, 759 760 powered prosthetic feet, 757 758 mechanical property tests, 758 759 musculoskeletal modeling and simulation, 759 prescription and expected use of, 750 752 activity bouts and durations, 751 activity in different environments, 751 752 activity levels, 751 Proteoglycans, 92 Proton density weighting (PD weighting), 301 Proximal intermetatarsal joints, 18 19 Proximal phalanges, 10 distal phalanges, 11 middle phalanges, 10 Pseudo tendon pathology, 305 PSS. See Post-polio syndrome (PSS) PT muscles. See Peroneus tertius muscles (PT muscles) PWB. See Partial weight bearing (PWB)

Q Quadratus plantae (QP), 33, 103 Quality-adjusted life years (QALYs), 743 Quantitative gait analysis, 79 Quantitative ultrasonography, 283

R Racing flats & spikes, 615 616 Radiation-free bone imaging, 318 319 Radiographs, 265 areas of future biomechanical research, 273 274 clinical X-ray measures of foot shape, 270 272 foot-specific applications and considerations, 270

Index

issues with X-ray measures of foot shape, 272 standard radiographic views of foot and ankle, 267 269 technology, 265 266 X-ray measurements of foot shape, 269 270 Radiography, 289 examination, 440 Radiostereometric analysis (RSA), 167, 169 171 ankle mortise width, 171 relationship between joints distal to talus, 170 talocrural joint, 169 technique, 169 transferral of rotation between leg and foot, 170 171 Randomized controlled trials (RCT), 628, 756 Range of motion (ROM), 507, 512 range of motion/flexibility/joint mobility, 440 441 Rearfoot strike (RFS), 107, 626 627 Reflective markers, 151, 153 Rehabilitation process, 492, 652 Relative stiffnesses, 416 Reliability assessment, 382 Repetition time (TR), 301 Residual hindfoot varus, 718 Residual reduction algorithm (RRA), 390 Residual torque depression (rTD), 222 Residual torque enhancement (rTE), 222 Resting calcaneal stance position (RCSP), 255 256 Retrocalcaneal bursitis, 520 Rheumatic diseases, 581 Rheumatic foot disease, 581 Rheumatoid arthritis (RA), 581 588, 642 643, 706 early rheumatoid arthritis, 582 established rheumatoid arthritis, 582 588 future research, 592 treatment goals in, 711 712 Rigid claw toes, 696 deformities, 692 Rigid flat foot, 482 Rigid hindfoot, 715 716 deformity, 672 Rizzoli MFM, 155 156 Robotic actuators, 357 Robotic Gait Simulator (RGS), 359 Rocker bottom shoes, 648 651, 656 design and prescription of ankle-foot orthosis, 652 654 design and prescription of rocker bottom shoes, 654 655 new designs, 655 patient populations, 649 652 ankle arthritis, 651 652 cerebral palsy, 651 limb salvage, 652 stroke, 650 651 roll-over shape, 649 sport applications, 655 656 variations on materials, 655

Rocker sole design, 654 Roll-over shape, 649 Root mean square (RMS), 172, 212 Rotation sequence, 71 Rubber waffle soles, 624 Running footwear, brief history of, 623 625 Running injury, 619 Running mechanics, 623 Running shoes, 615, 623 624 anatomy of, 612

S SACH. See Solid ankle cushion heel (SACH) Safety, ankle arthroplasty and arthrodesis, 741 742 Sag at talonavicular joint, 674 675 bony anatomy, 675 ligamentous failure, 675 677 surgical reconstruction, 677 679 Sagittal plane, 61 62 Scandinavian total ankle replacement (STAR), 734 735 Scarring, 436 Scars, 436 Screws, 716 717, 722 724 Segments, 152 153 Selective laser sintering, 655 Self-reported function and quality of life, 534 Sensitivity studies, 252 Sensors, 198 Sensory substitution, 236 237 Sensory testing, 443 Sequences with slice encoding for metal artifact correction (SEMAC), 311 Sequential anatomical rockers, 649 Sesamoiditis, 518 520 Shape modeling and assessment, 284 Shear stress, 666 Shear wave elastography, 324, 331, 333 Shock reduction, 140 Shod, 618 running, 618 Shoemakers, 701 Shoes innovations, 617 618 inserts, 637 types and features, 615 618 Short T2 tissues, 304 Short tau inversion recovery (STIR), 302 Silverskiold test, 692 Simulations, 387 388 Single-axis feet, 752 6-degrees-of-freedom (DOF), 352 Skeleton of foot, 3 5 Skin, 434 435, 598 marker artifact, 167 movement artifact in foot and ankle kinematics, 171 172 skin-mounted markers, 151 Small X-ray focal spot (SFS), 279 Soft tissues, 381, 692 injury, 381 reconstructive options, 710

775

Solescan Blanka, 342 Solid ankle cushion heel (SACH), 752 753, 760 Solid custom ankle-foot orthoses, 692 Solver selection, 381 382 Sonoelastography, 324 Spatiotemporal metrics, 755 756 Spinal cord injuries, 647 Splints, 539 Spondlyarthropathies, 589 Sports, 206 207 shoes, 616 617 Spring ligament, 703 704 reconstruction, 677 Sprinters, 107 Stability, 647, 649 metrics, 757 STAR. See Scandinavian total ankle replacement (STAR) Statics, 46 optimization, 390 simulators, 248 Statistical shape analysis techniques, 284 Step-cut “Z” osteotomy, 675 676 Step-cut osteotomy, 676 677 Stiffness, 49, 324 examining effects of, 756 757 Strain, 49, 324 elastography, 324 energy function, 415 416 Strength of materials, 48 49 Stress fractures, 316 317, 516 518 anatomical overview, 516 clinical impairments, 517 518 etiology, 516 517 mechanism of injury and pathomechanics, 516 517 treatment, 518 Stress radiographs, 708 Stress related disorders, 689 Strike pattern, 627 Strokes, 489 492, 647, 650 651 clinical treatment, 492 impact on foot function, 491 492 impact on kinematics, 490 pathology related to musculoskeletal system, 489 490 Stronger foot muscles, 631 Structural ex vivo testing, 140 142 Structural foot type, 452 454 Structural in vivo testing, 140 Structured light scanners, 342 Subchondral sclerosis, 682 Subtalar arthritis, 672 Subtalar arthrodesis, 711 712 Subtalar fracture/dislocation, 706 Subtalar fusion, 715 716 Subtalar impingement, 672 Subtalar joint (STJ), 13 14, 78, 124, 702 neutral position, 255 256 Superficial tangential zone (STZ), 92 Superficial veins, 42 Supramaximal tetanic contraction, 216 Surface electromyography, 212 213

776

Index

Surface scanning, 343 344 Surgical interventions, 374 Surgical techniques, brief description and history of, 732 735 Symptomatic flat foot, 643 Syndesmosis tears, 473 474 Syndesmotic view, 310 Systemic lupus erythematosus, 591

T Tailored magnetic resonance imaging protocol for foot and ankle, 308 311 imaging orientation, 308 metal artifact reduction sequences, 311 optimized imaging planes, 310 311 tailored magnetic resonance imaging protocols, 308 309 Talar fractures and dislocations, 706 treatment goals in, 712 Talar head, 702 Talar neck malunion, 687 Talocalcaneal angle, 708 Talocalcaneal joint, 96 97 Talocalcaneonavicular joint, 14 Talocrural joint, 78, 95 96, 169, 736 737 Talonavicular arthrodesis, 711 712 Talonavicular external rotation, 677 678 Talonavicular fusions, 714 716 Talonavicular joint (TN joint), 79, 97 98, 124, 552, 703, 717 sag at, 674 675 Talus, 6 fractures, 465 467 Taping, 540 Tarsal coalition, 480, 706 Tarsal tunnel syndrome, 498 Tarsometatarsal joints, 18, 99, 694 Tarsometatarsal (Lisfranc) injuries, 467 468 Tears, 304 Temperature, 435 drift, 200 Tendinopathy, 304, 514 516 anatomical overview, 515 clinical impairments, 516 etiology, 515 516 mechanism of injury and pathomechanics, 515 516 treatment, 516 Tendinosis, 304 Tendons, 113 114, 304, 313 315, 380 381, 442, 574, 597 pathology, 679 680 Tenosynovitis, 304 Therapeutic footwear, 662 665 Three-dimensional (3D) biometrics, 294 295 modeling, 375 376 motion analysis, 79 printed devices, 639 printed midsoles and outsoles, 618 printing, 638, 655 methods, 644 surface scanning of foot and ankle, 339

areas of future biomechanical research, 345 foot-specific applications and considerations, 342 345 history and development, 340 341 technologies, 341 342 tissue geometries, 250 Tibia, 3 Tibial nerves in foot, 40 41 Tibial plateau, 725 Tibialis anterior muscles (TA muscles), 21, 103 Tibialis posterior (TP), 27 28, 103 muscles, 103 tendon, 515 516 disease, 584 Tibiocalcaneal ligament, 420 Tibiofibular syndesmosis, 11 sprains, 403 Tibiospring, 123 Time to echo (TE), 301, 304 Time-of-flight systems, 342 Time-sequence MRI, 181 182 Toe deformities, 111 112 Toe separators, 539 “Toe walking” approach, 56, 494 495 biomechanical and musculoskeletal function, 494 495 diagnosis and etiology, 494 treatment, 495 Tornier Salto prosthesis, 741 742 Torque, 213 Total ankle arthroplasty (TAA), 100 designs, 734 Total ankle replacement (TAR), 733 734 Total contact cast (TCC), 662 Total joint replacements, 739 740 TR. See Repetition time (TR) Trabecular anisotropy, 92 Trabecular bone, 417 Transaxial plane. See Transverse plane Transfemoral amputation, 749 Transtibial amputation, 749 750 Transversal isotropy, 415 Transversally isotropic materials, 414 415 Transverse plane, 62 Transverse plane foot kinematics, 80 Transverse tarsal joint, 97 99 Traumatic foot and ankle injuries, 461 Achilles tendon rupture, 474 acute ankle sprains, 473 areas of future research, 475 calcaneal fractures, 464 465 metatarsal fractures, 468 472 midfoot crush injuries, 472 pilon fractures, 461 464 syndesmosis tears, 473 474 talus fractures, 465 467 tarsometatarsal (Lisfranc) injuries, 467 468 Triangulation systems, 342 Triceps surae, 25 26 Triphasic model, 94 Triple arthrodesis, 697 698, 711 712, 716 717

Triple-cut tendo-Achilles’ lengthening procedure, 681 Tube current modulation (TCM), 279 Tungsten microelectrode technique, 214 Twitch, 216 occlusion, 217 2D modeling, 375 376

U Ulcer, 566 diabetic footwear for ulcer prevention, 664 665 footwear and offloading for ulcer healing, 663 664 Ultrashort TE (UTE), 304 Ultrasonography (US), 708 710 Ultrasound, 142 143, 179, 532 533 assessment of Achilles tendon thickness in relation to weightbearing activities, 327 plantar fascia thickness in relation to weightbearing activities, 326 327 plantar soft tissue thickness in relation to weightbearing activities, 325 326 structural changes, 325 328 dynamometry, 330 elastography, 331 335 to assess mechanical properties of Achilles tendon, 333 334 to assess mechanical properties of plantar fascia, 333 to assess mechanical properties of plantar soft tissue, 331 332 limitations, 334 335 limitations of weightbearing ultrasound, 327 328 ultrasound assessment combined with measurement of load, 328 331 limitations, 331 ultrasound combined with dynamometry to assess mechanical properties of Achilles tendon, 329 331 ultrasound combined with load cells to assess mechanical properties of plantar soft tissue, 328 329 Ultrasound imaging, 323 works, 323 325 Unstable rocker bottom shoes, 654 655

V Validation, 252 253, 375, 390 Vantage, 735 Variable stiffness, 755 Variable-angle stiffness prosthetic foot, 754 755 Varus ankle arthritis, 699 Varus hindfoot alignment, 688 689 Veins, 42 deep veins, 42 superficial veins, 42 Ventral dorsal, 61 Vertical ground reaction force (vGRF), 759

Index

Video task analysis, 751 752 Virtual reality systems (VR systems), 236 Visco-hyperelastic model, 416 Viscoelastic mechanisms, 702 Viscoelasticity, 49 50 Volkmann’s triangle, 311 Volumetric, isotropic 3D magnetic resonance imaging sequences, 320

W Walking activity, 751 Wearable domains, 230 Wearable robots, 232 Wearable technologies, 229, 759 760 breakout box, 230 231 tech, 229 230

Weight bearing (WB), 255 256 Weight-bearing CT (WBCT), 289 290 advantages and limitations, 296 biases of conventional radiography, 290 291 future areas of research, 296 297 indications, 294 scans, 671, 675, 677 678, 691 technical aspects, 291 294 3D biometrics, 294 295 Weight-bearing radiographs, 690 691 Weightbearing ankle radiographs, 671

777

clinical X-ray measures of foot shape, 270 272 generators, 182 183 imaging, 180, 452 issues with X-ray measures of foot shape, 272 markers, inserting, 175 measurements of foot shape, 269 270 stereophotogrammetry, 181 182

Y Yield/ultimate/failure load and strain, 49

X

Z

X-rays, 265, 690 691

Zener model, 416 417