BIOMECHANICAL INVESTIGATIONS AND COMPUTATIONAL
MODELING OF THE HUMAN ANKLE
By

Feng Wei

A DISSERTATION

Submitted to
Michigan State University
in partial fulfillment of the requirements
for the degree of

DOCTOR OF PHILOSOPHY

Engineering Mechanics

2012

ABSTRACT
BIOMECHANICAL INVESTIGATIONS AND COMPUTATIONAL MODELING OF
THE HUMAN ANKLE
By
Feng Wei
The ankle is one of the most common sites for acute musculoskeletal injuries, and sprains
account for 75% of ankle injuries. Acute ankle trauma is responsible for 10 to 30% of
sports-related injuries in young athletes. Each year, an estimated one million persons
present to physicians with acute ankle injuries. More than 40% of ankle sprains have the
potential to cause chronic problems. While lateral ankle sprains are the most common
injury, high ankle sprains represent a more disabling problem and require a longer
recovery period and different treatment. Injury to the tibiofibular syndesmosis ligaments,
which bind together the distal ends of the tibia and fibula, is commonly referred to as a
high ankle sprain. External rotation of the foot has been implicated in high ankle sprains,
but the mechanism of injury is still unclear and the pathologies due to external foot
rotation remain controversial. The mechanisms of lateral ankle sprains are better
understood to evolve from foot inversion, but the level of strain being produced in lateral
ligaments during such injuries is yet unclear.

The mechanisms of sports-related injuries can be studied with various approaches:
cadaver experiments, kinematic studies and computational modeling. The advantage of
human cadaver tests is that the forces and motions experienced by the ankle joint can be
controlled and measured with great precision. Dissections of joints after testing can also
allow for an in-depth study of ligament and bone injuries. While ligament damages due to

excessive strains are often involved in most sports-related injuries, studies with cadaver
ankles may not reflect ligament behaviors during injury-producing events in living
humans. Three-dimensional simulations of human movements using dynamic,
computational models can offer an attractive method to separate various motions of the
ankle bones, rapidly solve for motion-based mechanics, and determine ligament strains
during physiological motions.

In this dissertation, cadaver studies were conducted to investigate ankle responses to
external foot rotation with different foot and shoe constraints. A computational ankle
model has also been developed and validated against cadaver experiments and studies
with human subjects. These studies showed that external rotation of a highly everted foot
generated a high ankle sprain (ligamentous damage to the anterior tibiofibular ligament),
while external rotation of a neutral foot produced a ligamentous injury to the anteromedial aspect of the ankle (damage to the anterior deltoid ligament). In addition to the
failure level testing, a low level testing with football shoes showed that while flexible
shoes generated lower joint torques than rigid shoes during external foot rotation, they
produced more talus eversion that may induce strain in the anterior tibiofibular ligament.
Ligamentous injuries due to excessive foot inversion have also been studied using the
computational model in this dissertation. These studies provide a better understanding to
the mechanisms of various ankle injuries that might aid in prevention programs and
treatment strategies for these trauma patients.

ACKNOWLEDGEMENTS

I would like to thank my mentor Dr. Roger Haut for his expertise, leadership, support,
and dedication throughout my research at the Orthopaedic Biomechanics Laboratories
(OBL). I would also like to acknowledge Drs. Seungik Baek, Dahsin Liu, and John
Powell for their insightful advice and serving on my committee. In addition, I would like
to thank Mr. Clifford Beckett for technical assistance, Dr. Joel Post for clinical advice,
and Mrs. Jean Atkinson for help in specimen preparation during my study at the OBL.
Finally, I would like to acknowledge everyone who worked with me at the OBL for their
help and friendship: Eric Meyer, Jerrod Braman, Nurit Golenberg, Brian Weaver, Keith
Button, Mark Villwock, Daniel Isaac, Tim Baumer, Brian Powell, and Kaitlyn Kopke.

iv

TABLE OF CONTENTS

LIST OF TABLES ............................................................................................................ vii
LIST OF FIGURES ......................................................................................................... viii
CHAPTER 1: Introduction and background.
Ankle kinematics ·····································································································1
Anatomy of the distal tibiofibular syndesmosis·······················································3
Review of experimental literature ············································································7
Review of computational literature ········································································12
Summary and objectives ························································································14
References ··············································································································16
CHAPTER 2: A biomechanical investigation of ankle injury under excessive external
foot rotation in the human cadaver.
Abstract ··················································································································21
Introduction ············································································································22
Methods··················································································································23
Results ····················································································································27
Discussion ··············································································································32
References ··············································································································36
CHAPTER 3: Development and validation of a computational model to study the effect
of foot constraint on ankle injury due to external rotation.
Abstract ··················································································································40
Introduction ············································································································41
Methods··················································································································44
Results ····················································································································54
Discussion ··············································································································60
References ··············································································································65
CHAPTER 4: Simulation studies to investigate ankle sprain mechanisms.
Part I: Mechanism of injury in a lateral ankle sprain.
Introduction ············································································································70
Methods··················································································································71
Results ····················································································································74
Discussion ··············································································································75
Part II: Mechanism of injury in a high ankle sprain.
Introduction ············································································································77
Methods··················································································································78
Results ····················································································································80
Discussion ··············································································································81
References ··············································································································83
v

CHAPTER 5: Eversion during external rotation of the human cadaver foot produces high
ankle sprains.
Abstract ··················································································································85
Introduction ············································································································86
Methods··················································································································88
Results ····················································································································93
Discussion ··············································································································98
References ············································································································106
CHAPTER 6: Rotational stiffness of football shoes influences talus motion during
external rotation of the foot.
Abstract ················································································································110
Introduction ··········································································································112
Methods················································································································114
Results ··················································································································120
Discussion ············································································································126
References ············································································································132
CHAPTER 7: A computational model to investigate shoe and shoe-surface interface
effects on ankle ligament strains during a simulated sidestep cutting task.
Abstract ················································································································136
Introduction ··········································································································137
Methods················································································································137
Results ··················································································································139
Discussion ············································································································141
Conclusion ···········································································································143
References ············································································································145
CHAPTER 8: Determination of dynamic ankle ligament strains from a computational
model driven by motion analysis based kinematic data.
Abstract ················································································································147
Introduction ··········································································································149
Methods················································································································150
Results ··················································································································155
Discussion ············································································································162
References ············································································································166
CHAPTER 9: Conclusion. ·······························································································172
APPENDIX ······················································································································188

vi

LIST OF TABLES

Table 2-1: Torque and foot rotation at failure ...................................................................30
Table 3-1: Ligament abbreviations and stiffnesses ............................................................46
Table 3-2: Sensitivity analysis results................................................................................57
Table 5-1: Specimen descriptions and test results .............................................................96
Table 5-2: Talus motion relative to the tibia ......................................................................97
Table 6-1: Specimen descriptions and test results ...........................................................123
Table 6-2: Talus motion in different shoes ......................................................................124
Table 8-1: Variance of trials from each subject ...............................................................155
Table 9-1: Torque and rotation at failure .........................................................................174
Table 9-2: Foot rotation at failure and resultant injury ....................................................180
Table A-1: Raw data from Chapter 5 ...............................................................................197
Table A-2: Raw data from Chapter 6 ...............................................................................201
Table A-3: SolidWorks input and output from Chapter 8 ...............................................203

vii

LIST OF FIGURES

Figure 1-1: Bony anatomy of the hindfoot and ankle ..........................................................1
Figure 1-2: Terminology describing motion of the foot ......................................................2
Figure 1-3: Ligaments of the distal tibiofibular syndesmosis ..............................................4
Figure 1-4: Medial view of osteoligamentous ankle anatomy .............................................6
Figure 1-5: Lateral view of osteoligamentous ankle anatomy .............................................7
Figure 2-1: Screw positioning and foot potting .................................................................24
Figure 2-2: Experimental setup ..........................................................................................25
Figure 2-3: Torsional stiffness of foot/ankle complex .......................................................28
Figure 2-4: Torque-angular displacement response ...........................................................28
Figure 2-5: Failure modes resulting from external rotation ...............................................31
Figure 3-1: Reconstructed 3D model of the foot anatomy ................................................45
Figure 3-2: Various models used in the study ...................................................................49
Figure 3-3: Schematic of the foot with a screwed marker set............................................52
Figure 3-4: Comparison between model and experiment ..................................................53
Figure 3-5: Comparison of torque-rotation behavior .........................................................54
Figure 3-6: Max torque and rotational stiffness .................................................................55
Figure 3-7: Strains in different ligaments of the potted foot .............................................56
Figure 3-8: Strains in ligaments during external rotation of the foot ................................58
Figure 3-9: Talus rotation relative to the tibia ...................................................................60
Figure 4-1: Schematic of the generic ankle model ............................................................73
Figure 4-2: Ankle model in various positions....................................................................74
viii

Figure 4-3: Strains in various ligaments for different motions ..........................................75
Figure 4-4: Ankle model in various positions....................................................................80
Figure 4-5: Strains in various ligaments for different motions ..........................................81
Figure 5-1: Specimen preparation and setup .....................................................................90
Figure 5-2: Testing setup ...................................................................................................92
Figure 5-3: Torque-rotation curves of different limbs .......................................................94
Figure 5-4: Typical temporal profile and torque-rotation curves ......................................95
Figure 5-5: Typical gross injuries ......................................................................................97
Figure 5-6: Schematics of the ankle in anterior and medial views ..................................100
Figure 6-1: Shoe stiffness tests preparation and setup .....................................................115
Figure 6-2: Cadaver tests preparation and setup ..............................................................117
Figure 6-3: Testing setup .................................................................................................119
Figure 6-4: Torque-rotation curves from shoe stiffness tests ..........................................120
Figure 6-5: Rotational stiffness of torque-rotation curves ...............................................121
Figure 6-6: Torque-rotation curves of cadaver tests ........................................................122
Figure 6-7: Comparisons of temporal profiles of talus rotation ......................................125
Figure 6-8: Ankle ligament strains from a computational model ....................................126
Figure 7-1: Foot and surface models ...............................................................................139
Figure 7-2: Strains in various ligaments ..........................................................................141
Figure 7-3: Strains in various ligaments during external rotation ...................................141
Figure 8-1: Marker set of the Oxford Foot Model ...........................................................151
Figure 8-2: Generic ankle model showing the simulated ligaments ................................154
Figure 8-3: Typical profiles of moment and rotation.......................................................156

ix

Figure 8-4: Hindfoot kinematics and kinetics in three orientations .................................157
Figure 8-5: Max strains in various ligaments ..................................................................158
Figure 8-6: Temporal profiles of hindfoot motions and ligament strains ........................160
Figure 8-7: Temporal sequences of peak motions and peak strains ................................161
Figure 9-1: Strains in different ankle ligaments ..............................................................175
Figure 9-2: Ankle ligament strains for different motion..................................................177
Figure 9-3: Ankle ligament strains under external rotation .............................................178
Figure 9-4: Ankle ligament strains in different shoes ......................................................181
Figure 9-5: Ankle ligament strains for different surfaces ................................................183
Figure 9-6: Ankle ligament strains for in vivo tests .........................................................184
Figure A-1: A typical MIMICS interface ........................................................................191
Figure A-2: Calculations of foot rotation angles in a JCS ...............................................195

x

CHAPTER 1
INTRODUCTION AND BACKGROUND

Ankle kinematics
The ankle plays a fundamental role in
locomotion. The commonly called ankle
actually consists of three joints: the
distal tibiofibular joint, the tibiotalar
joint or talocrural joint, often referred to
as the ankle joint, and the talocalcaneal
joint, often referred to as the subtalar
joint (Figure 1-1). Ankle motion is
typically

described

based

on

the

rotational angle of either the calcaneus
or the forefoot relative to the tibia about Figure 1-1 Bony anatomy of the hindfoot and
three Cartesian axes (Figure 1-2). These ankle.
axes are in the anterior-posterior (x-axis),
medial-lateral (y-axis), and inferior-superior (z-axis) directions. Unfortunately, the
terminology describing the rotational motion of the foot-ankle complex is somewhat
varied in the literature. Rotation about the medial-lateral axis (y-axis) is almost invariably
called dorsiflexion (toes towards the leg) or plantarflexion (toes away from the leg).
Rotation about the long axis of the foot (x-axis) is typically called inversion (sole of the
foot facing the midline) or eversion (sole of the foot facing away from the midline). The

1

terms adduction and supination are generally used synonymously with inversion, and the
terms abduction and pronation are generally used synonymously with eversion, although
sometimes supination and pronation are used to describe multiaxial forefoot rotation and
adduction and abduction are used to describe uniaxial hindfoot rotation [18]. Rotation of
the foot about the long axis of the leg (z-axis) is typically called internal rotation (toes
point towards the midline) or external rotation (toes point away from the midline).

Figure 1-2 Terminology used to describe the motion of the foot with respect to the leg.

Normal unconstrained motion of the foot/ankle complex does not occur about a fixed
Cartesian axis. Rather, due to the bony anatomy and ligamentous constraints of the ankle
and subtalar joints, the foot tends to rotate about a combined and moving axis. Ankle
joint motion is primarily responsible for dorsiflexion and plantarflexion, whereas subtalar
joint motion is primarily responsible for inversion and eversion. Internal and external
rotations involve roughly equal contributions from the ankle and subtalar joints. Inversion
and internal rotation are strongly coupled to each other and weakly coupled to
2

plantarflexion. Likewise, eversion and external rotation are strongly coupled to each
other and weakly coupled to dorsiflexion. Forced rotation about any one of these axes
induces rotation about the other two coupled axes during normal, unconstrained motion
of the foot/ankle complex [22,26,32,35]. Injury, however, typically occurs because the
motion of the foot relative to the leg is constrained and abnormal.

Anatomy of the distal tibiofibular syndesmosis
The connection between tibia and fibula is formed by three structures, namely the
superior tibiofibular joint, the interosseous membrane, and the distal tibiofibular joint.
The proximal and distal tibiofibular joints are syndesmoses, which are fibrous joints with
ample intervening fibrous connective tissue. The distal tibiofibular syndesmosis consists
of four ligaments (Figure 1-3): the anterior inferior tibiofibular ligament, the posterior
inferior tibiofibular ligament, the transverse ligament and the interosseous ligament,
which is the most distal part of the interosseous membrane [1,4].

The anterior inferior tibiofibular ligament (ATiFL)
The ATiFL is a strong, shiny ligament, maximal 2 cm in width and almost 0.5 cm thick.
It consists of 3 bundles, separated by 2 mm wide gaps that slightly converge in the laterodistal direction [1]. The superficial anterior fibers are 2 to 3 cm long, the deeper posterior
fibers somewhat shorter [6]. They run obliquely downward from the anterior tibial
tubercle to the antero-medial aspect of the fibular malleolus with an angle of about 30-50
degrees.

3

Figure 1-3 Anterior, posterior, and lateral views of select ligaments of the distal tibiofibular syndesmosis: the anterior-inferior
tibiofibular ligament (ATiFL); the posterior-inferior tibiofibular ligament (PTiFL), of which the inferior transverse ligament (ITL) is
part; and the interosseous ligament (IOL), which represents the thickened distal part of the interosseous membrane. The arrows
indicate the respective location and point to the cross-sectional view [20].

4

The posterior inferior tibiofibular (PTiFL) and the transverse (ITL) ligament
The PTiFL has a trapezoid shape. It is more compact and runs more horizontal than the
ATiFL. It is on average 18 mm width and 0.6 mm thick. Superiorly there is an almost
continuous transition into the interosseous membrane [1]. The distal margin of the PTiFL
is formed by a more anterior and transverse bundle, which is recognized by some authors
as a separate entity known as the transverse ligament [6], and serves as a labrum to the
tibia [44].

The interosseous ligament (IOL)
The tibiofibular interosseous ligament is a strong pyramid shaped thickening of the fibers
of the interosseous membrane. These fibers form a network that runs from 5 to 1.5 cm
proximal to the tibiotalar joint space in a latero-distal direction from the tibia [1]. The
posterior edge of the IOL almost continuously passes into the PTiFL. The anterior surface
of IOL is divided from the ATiFL by a small gap.

The deltoid ligament
Although the deltoid ligament is not part of the distal tibiofibular syndesmosis it is
functionally closely associated with it. The deltoid ligament originates from the medial
malleolus and is covered posteriorly by the tendons of tibialis posterior and flexor
digitorum longus. The deltoid ligament consists of superficial and deep layers. In general,
the superficial layer originates primarily from the distal tibia and inserts on the navicular
bone, the spring (plantar calcaneonavicular) ligament, and the medial tubercle of the talus,
whereas the deep layer runs from the intercollicular groove and the posterior colliculus to

5

the medial surface of the talus [4]. While it is difficult to distinguish the bundles of the
deltoid ligament, many authors refer to them as four ligaments: the anterior and posterior
tibiotalar ligaments (ATTL and PTTL), tibionavicular ligament (TiNL), and
tibiocalcaneal ligament (TiC) (Figure 1-4). The ATTL and TiNL have an antero-distal
direction and often combined and referred to as the anterior deltoid ligament.
Comparably, the PTTL and TiCL are oriented postero-distally often referred to as the
posterior deltoid ligament.

Figure 1-4 Medial view of the osteoligamentous ankle anatomy.

6

Review of experimental literature
Inversion related ankle injuries
Modern

biomechanical

cadaveric

studies have confirmed the accuracy
of the Lauge-Hansen scheme with
regard

to

injuries

and

supination-adduction
have

contributed

additional detail. By far the most
common site of inversion injury is
the lateral ligamentous complex
(Figure 1-5). Dias [15] reported that
the first stage of inversion injury is
the anterior talofibular ligament Figure 1-5 Lateral view of the osteoligamentous
(ATaFL)

if

plantarflexion

the

ankle
and

is

in ankle anatomy.
the

calcaneofibular ligament (CFL) if the ankle is in neutral flexion. With the ankle in neutral
flexion, some studies have also reported primarily CFL injuries [36,37], whereas others
have reported a roughly even distribution between CFL and ATaFL injuries [2,19,39,].
Additional ligaments, most notably the posterior talofibular ligament (PTaFL), are
injured if the ankle continues to invert [15,39]. A similar injury pattern was reported for
loading about a combined dorsiflexion-inversion axis [39]. Inversion of a dorsiflexed
ankle has produced injury of the CFL and other ligaments in three specimens, one of
which also sustained a fracture of the sustentaculum tali in the calcaneus [19]. An

7

accidental Grade I ATaFL sprain sustained in a human volunteer in a laboratory study
was documented to be the result of combined inversion and internal rotation with the
ankle in 15° of dorsiflexion [17]. This result suggests that the ATaFL can be injured
without the foot in plantarflexion if inversion is coupled with internal rotation.

Without axial compression of the ankle joint, lateral malleolar fractures have been
reported only rarely in inversion [15,19,36,39]. Osteochondral fractures of the ankle joint
due to inversion have been reported with axial compression but not without it [19]. With
axial compression of one to three times body weight, lateral malleolar fractures and
avulsions occur more frequently in inversion, but lateral ligament injury remains the
predominant failure mode [2,19]. Specimens sustaining fractures and avulsions in
inversion have been noted to have lower bone mineral densities than specimens
sustaining midsubstance ligament tears [19].

Compression of the leg onto a wedge that slopes down laterally at a sufficiently steep
angle induces unstable inversion of the ankle, and therefore is considered here to be a test
of inversion as injury mechanism rather than axial compression. Self et al. [42] and
Konradsen and Voigt [27] produced ATaFL tears initially, followed by CFL tears in this
mode of loading with a 30° wedge, whereas Hirsch and Lewis [24] produced mostly
lateral and bimalleolar fractures. The discrepancy may be related to the inversion angle
and method of foot fixation, which were not described by Hirsch and Lewis [24].

8

External rotation related ankle injuries
External rotation injuries of the ankle have been studied extensively, particularly
supination-external rotation (SE). These studies have produced some conflicting results
and have generally failed to support the Lauge-Hansen scheme. When Dias and Foerster
[16] attempted to reproduce the results of Lauge-Hansen [30] by applying external
rotation to maximally inverted ankles without axial loading, they initially produced total
ligamentous and capsular tears with the ATiFL remaining intact. They considered these
six tests to be failures. In 8 out of 10 subsequent tests, they artificially cut the ATiFL
(SE1) and subsequently produced SE2, SE3, and SE4 injuries. Dias and Foerster [16]
considered a tear of the ATiFL to be a prerequisite for the occurrence of the subsequent
stages of SE injury. Markolf et al. [33], using the same general methods, obtained short
oblique fibular fractures in two thirds of their specimens and lateral ligament tears in the
others, all without ATiFL injury.

Hirsch and Lewis [24], also using the same general methods but with the addition of body
weight, generated lateral malleolar fractures in all six of their injured specimens,
syndesmosis injury in two specimens, as well as three medial malleolar fractures and two
deltoid ligament tears. Stiehl et al. [43], using the same loading configuration, reported
that an axial load level of at least three times body weight was required to consistently
produce fracture. They produced oblique fibular fractures with torn ATiFL and deltoid
ligaments in two thirds of their male specimens and isolated transverse fibular fractures
in almost half of their female specimens. Michelson et al. [34], using generally the same
methods as Hirsch and Lewis [24], produced low fibular fractures in only two out of 10

9

specimens and lateral ligament avulsions in five others. The ratio of lateral ligament
avulsion to fibular fracture increased when the ankle was externally rotated while being
held in inversion and plantarflexion [34].

Externally rotating the ankle while holding the foot in dorsiflexion and inversion
generated short oblique fibular fractures with postero-superior fracture lines in all 24 of
Schaffer and Manoli's [41] specimens with only two associated ATiFL injuries. Similar
results were obtained by Michelson et al. [34] and Wei et al. [45] in this loading
configuration. Completely different results were obtained by Boon et al. [5], who
reported generating six isolated LPT fractures by external rotation of dorsiflexed (20°)
and inverted (10°) ankles. However, Boon et al. [5] conducted their tests with a much
higher level of superimposed axial compression (2200-8900 N) than did Schaffer and
Manoli [41] (147 N), Michelson et al. [34] (700 N), or Wei et al. [45] (2000 N).

Michelson et al. [34] reported that adding a “valgus load” to the ankle moved the site of
fibular fracture from below the tibial plafond up to the level of the tibial plafond. This
was accomplished by initially positioning the leg to a position of 6–8° of valgus before
testing, which would seem to be equivalent to increasing the inversion angle of the ankle
from 25° to 31–33° and adding a small component of plantarflexion to the external
rotation torque. Michelson et al. [34] reported that positioning the foot in dorsiflexion (in
addition to the inversion, body weight compression, and valgus leg position) before
externally rotating it resulted in all five specimens sustaining a transverse fracture of the
medial malleolus in addition to fibular fractures at the level of the tibial plafond, where

10

other loading conditions had resulted in either no medial injury or only the occasional
deltoid ligament tear.

In spite of the fact that the ATiFL is supposed to be the site of initial injury in SE loading
according to the Lauge-Hansen classification, only one experimental study [43] has
reported a high incidence of ATiFL tears with bone fractures in this loading mode. The
remaining studies have all reported ATiFL injuries in about one quarter or fewer
specimens tested in SE loading [5,16,24,33,34,41,45]. Likewise, only one experimental
study has reported a high incidence of PTiFL tear or posterior malleolar fracture in SE
loading [16]. According to the Lauge-Hansen classification, injury to the posterior
structures should precede injury to the medial structures. However, Michelson et al. [34]
reported PTiFL injury in only 4 of 10 specimens with medial injury, and Stiehl et al. [43]
reported no PTiFL injury in 21 specimens with medial injury.

A recent study by Haraguchi and Armiger [21] has shown that external rotation of an
everted foot axially loaded to body weight (700 N) typically results in the classic LaugeHansen SE injury pattern: a short oblique fibular fracture starting at the level of the tibial
plafond and running in a postero-superior direction, disruption of the ATiFL and PTiFL
complexes, and either a medial malleolar fracture or a deltoid ligament tear. They noted
that the lateral malleolar fracture preceded medial injury and that PTiFL tear or posterior
malleolar fracture could occur before or after medial injury. This study provides a more
logically consistent injury mechanism to account for the large number of SE fractures
seen clinically. Interestingly, this finding has been confirmed by an analysis of injury

11

videos posted on YouTube.com in which five out seven ankles subjected to a pronationexternal rotation injury mechanism displayed a radiographic fracture pattern that would
be classified as supination-external rotation under the Lauge-Hansen scheme [29].

Haragurchi and Armiger [21] also noted that adding a 100 N lateral shear force to the
bottom of the foot resulted in similar injury patterns, but with a significantly higher
incidence (three out of eight specimens) of high fibular fractures with a reversed
(postero-inferior) or comminuted fractured line, which is the classic Lauge-Hansen PE
fracture pattern. Fibular fractures occurred below the level of the tibial plafond in the
minority of specimens in which the ATiFL complex remained intact, which is in
agreement with the findings of Stiehl et al. [43]. Hirsch and Lewis [24], using similar
methods, reported results similar to Haraguchi and Armiger [21], though in less detail.
Experimental studies have reported that for neutrally flexed feet, external rotation and
rotation about a combined external rotation/dorsiflexion axis result in a fairly even
distribution of medial and lateral injuries with a high incidence of fractures and a low
incidence of syndesmosis injuries [2,24,39].

Review of computational literature
Computational models of musculoskeletal joints and limbs can provide useful
information about joint interactions, joint kinematics, spatial

positioning of

bones/appendages, muscle/ligament lengths, muscle forces and moments, joint contact
pressures, and range of motion predictions [31]. These validated models can also be used
for simulation of injuries and repairs, effects of fractures/osteotomies, surgical

12

reconstructive procedures, changes in joint angles, changes due to ligament
lengthening/shortening or deficiencies, and variations due to different tendon attachment
sites. Two common approaches for the development of these biomechanical models are
finite element analysis based and multibody kinematic or dynamic based [714,25,23,28,38,40,46]. The one chosen typically depends on the information sought as
both approaches have their advantages and yield much useful information. A primary
advantage of multibody modeling is its ability to solve for mechanics of large structures
using highly efficient algorithms, which can execute much faster than the continuumbased finite element analysis (FEA) [28]. Multibody modeling has been applied to the
upper and lower extremities as well as to specific segments such as the knee and
wrist/hand complex. Many of these multibody models depend on experimental
measurements of bone movements, assumptions about joint degrees of freedom, and
moments associated with muscle forces to guide the motion in the computational models
[7-9,12-14,25,38]. Fewer models incorporate articulation geometry, ligaments, and other
anatomical features to guide joint function. Often this is accomplished by coding of
governing mechanics equations and solution algorithms into custom computer programs.
For knee kinematics using rigid three-dimensional multiple bodies and springs to
represent ligaments, Wismans et al. [46] and Hirokawa [23] performed studies using rigid
surface contacts between the bones of the knee joint while Kwak et al. [28] allowed a
small amount of overlap to account for articular cartilage.

13

Summary and objectives
Sports-related injury mechanisms can be studied through various approaches: cadaveric
experiments, kinematics, biomechanics, and computational modeling. The great
advantage of human cadaveric testing is that the forces and motions experienced by the
cadavers during testing can be controlled and measured with great precision. Dissection
of cadaveric specimens after testing allows for a more in-depth study of injurious
pathology than typically can be gained through clinical testing or imaging. While
ligament damage due to excessive strains is often involved in sports-related injuries,
cadaveric studies, however, may not reflect ligament behaviors during injury-causing
events. Dynamic three-dimensional simulation studies offer an attractive alternative
because of their ability to separate motions and study them individually, rapidly solve for
motion-based mechanics, and determine ligament strains during large motions.

The ankle is one of the most common sites for acute musculoskeletal injuries, and sprains
account for 75% of ankle injuries. Acute ankle trauma is responsible for 10 to 30% of
sports-related injuries in young athletes. Each year, estimated one million persons present
to physicians with acute ankle injuries. More than 40% of ankle sprains have the potential
to cause chronic problems. Injury to the tibiofibular syndesmosis ligaments, which bind
together the distal ends of the tibia and fibula, is commonly referred to as a high ankle
sprain. While foot inversion involved lateral ankle sprains are the most common injury,
high ankle sprains represent a more disabling problem and require a longer recovery
period and different treatment. External rotation of the foot has been implicated in high

14

ankle sprains, but the mechanism of injury is still unclear and the pathologies due to
external foot rotation remain controversial.

In this dissertation, series cadaveric experiments were conducted to investigate the ankle
responses to various levels of external foot rotation under different foot constraint
conditions. In addition, a computational ankle model was developed and validated against
cadaveric experiments and gait analyses. The mechanism of injury in a high ankle sprain
was examined using both approaches. These studies may provide a valuable basis for
understanding, prevention, and treatment of ankle injuries.

15

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16

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Potier P, Petit P. 2002. The effects of axial preload and dorsiflexion on the
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Journal of Bone and Joint Surgery (American Volume) 69:596-604.
42. Self BP, Harris S, Greenwald RM. 2000. Ankle biomechanics during impact
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43. Stiehl JB, Skrade DA, Johnson RP. 1992. Experimentally produced ankle
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44. Stoller D. 1998. MRI, arthroscopy, and surgical anatomy of the joints. Lippincott
Williams and Wilkins, Philadelphia.
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20

CHAPTER 2
A BIOMECHANICAL INVESTIGATION OF ANKLE INJURY UNDER EXCESSIVE
EXTERNAL FOOT ROTATION IN THE HUMAN CADAVER

ABSTRACT:
Background: Numerous studies on the mechanisms of ankle injury deal with injuries to
the syndesmosis and anterior ligamentous structures, but a previous sectioning study also
describes the important role of the posterior talofibular ligament (PTaFL) in the ankle’s
resistance to external rotation of the foot. It was hypothesized that failure level external
rotation of the foot would lead to injury of the PTaFL.
Method of approach: Ten ankles were tested by externally rotating the foot until gross
injury. Two different frequencies of rotation were used in this study, 0.5 Hz and 2 Hz.
Results: The mean failure torque of the ankles was 69.5 ± 11.7 Nm with a mean failure
angle of 40.7 ± 7.3 degrees. No effects of rotation frequency or flexion angle were noted.
The most commonly injured structure was the PTaFL. Visible damage to the syndesmosis
only occurred in combination with fibular fracture in these experiments.
Conclusions: The constraint of the subtalar joint in the current study may have affected
the mechanics of the foot and led to the resultant strain in the PTaFL. In the real-world,
talus rotations may be affected by athletic footwear that may influence the location and
potential for an ankle injury under external rotation of the foot.

Keywords: posterior talofibular ligament, anterior deltoid ligament, ankle sprain,
external rotation, sports injury

21

INTRODUCTION
In the National Football League (NFL), excessive external rotation of the foot is often
associated with significant time loss injuries [1,2]. These injuries typically involve the
ligamentous structures of the ankle joint. The importance of individual ligamentous
contributions to resistive torque during external rotation of the foot has been investigated
in numerous sectioning studies [3-5]. This involves surgically cutting a ligament and
quantifying a joint motion. Buckle transducers and strain gauges have also been used to
show that there are relatively high forces developed in the ankle ligaments during
external rotation of the foot [6,7]. Yet, in contrast to the soft tissue injuries of the ankle
joint experienced during sports play [1,8,9], experimental studies have typically
generated a high frequency of bone fractures when the foot is externally rotated to failure
[10-12].

In the current literature there are conflicting reports as to the specific ligaments that
provide the primary restraint to external rotation of the foot. Numerous studies on the
mechanisms of ankle injury deal with injuries to the syndesmosis and anterior
ligamentous structures [11,13,14]. However, a previous laboratory study using sectioning
techniques describes the important role of the posterior talofibular ligament (PTaFL) in
the ankle’s resistance to physiological levels of external rotation of the foot [15]. In
addition, Colville et al. [16] measured strain in various ankle ligaments and concluded
that the greatest strain experienced during dorsiflexion (20°) and external rotation is in
the PTaFL. And, Rasmussen [17] showed that the PTaFL restricted dorsiflexion of the
foot. Clinicians also indicate a high frequency of cases involving injury to this posterior

22

ligamentous structure [16,18]. Yet some clinical studies report that the PTaFL is rarely
injured, except in association with complete dislocation of the talus [19].

While studies have shown an important relation between external rotation of the foot and
physiological levels of strain in the PTaFL, the objective of the current study on the
human cadaver ankle was to document failure characteristics of this joint under
controlled external rotation of the foot. The hypothesis was that an excessive external
rotation of the foot would lead to isolated injury of the PTaFL. The data may be helpful
in the development of future strategies for injury diagnosis and prevention. The
biomechanical data may also be used to develop a surrogate ankle that mimics the human
response that could be used to evaluate the athletic shoe-surface interfaces for studies of
injury risk to the ankle [20].

METHODS
Torsion experiments were conducted on ten lower limbs from male cadavers (aged 43
17 years). The limbs were procured through university sources (see Acknowledgements),
stored at -20 degrees Celsius and thawed to room temperature for 24 hours prior to
testing. The tibia and fibula were transected approximately 15 cm distal to the center of
the knee. The proximal end of the tibia and fibula shafts were cleaned with 70% alcohol
and potted in a rectangular aluminum tray with room temperature curing epoxy (Fibre
Strand, Martin Senior Corp., Cleveland, OH). Several screws were placed from medial
and lateral aspects into the calcaneus with approximately 15-20 mm of the length
projected out so as to constrain the calcaneus in the potting material (Figure 2-1). The

23

foot was surrounded and supported with epoxy. Care was taken to leave space around the
medial and lateral malleoli (Figure 2-2A). Various bones had motion sensing markers
inserted via screws, but the motion data was limited and not reported here. The foot was
positioned in slightly different angles of flexion (dorsi- or plantar-) in these experiments
by means of interchangeable wedges placed underneath the foot plate (Figure 2-2B). The
foot was also everted (10 to 15 degrees about an anteroposterior axis) in these
experiments by adjustment of the eversion fixture (Figure 2-2B). The flexion and
eversion varied in an attempt to obtain a midsubstance ligament rupture as to the
literature [16,17].

A

B

C

Figure 2-1 Screw positioning from medial (A) and lateral (B) sides, and foot potting (C).

24

A
B
Vertical
Linear Actuator

Eversion
Fixture

Plate Allowing
X-Y Adjustment

Dorsiflexion
Wedge

Load
Cell
Rotary
Actuator
Figure 2-2 (A): Distal lower extremity constraints. Foot fixated in potting material with reflective marker arrays attached to bony
landmarks. This marker set was used for a pilot motion analysis study (not documented here). (B): Schematic of experimental setup.

25

A custom, hydraulic, biaxial testing machine was built by mounting a 244 Nm rotary
actuator (Model SS-001-1V, Micromatic, Berne IN) onto a linear actuator frame (Model
312.21, MTS Corp., Eden Prairie, MN) with a vertically oriented linear actuator (Model
204.52, MTS Corp.). The actuators had separate controllers (Model 458.2 Microconsole
for the linear actuator and Model 442 Controller for the rotary actuator, MTS Corp.). The
rotational displacement was programmed by a waveform generator (Model 458.91
Microprofiler, MTS Corp.) that produced a haversine waveform. The foot was attached to
the rotary actuator through a biaxial (torsion-axial) load cell (Model 1216CEW-2K,
Interface, Scottsdale, AZ). The axis of rotation of the foot was aligned with the ankle
center between the medial and lateral malleoli. The proximal end of the extremity was
secured to the rotation locked linear actuator with a custom fixture designed to allow
sufficient positioning for flexion and eversion of the foot. Each specimen was mounted
with the tibia axis aligned with the linear actuator.

Compressive preloads of 2000 N (approximately two to three times body weight) were
applied through the tibia prior to the application of an external rotation. A dynamic
angular rotation was input with a position controlled waveform. Two different
frequencies of rotation were used in this study, 0.5 Hz (1 sec to peak) and 2 Hz (0.25 sec
to peak). The rotation was increased by increments of five degrees in successive tests
until diagnosed ligamentous injury or bone fracture. Approximately 5-10 minutes were
allowed between increments for the researchers to observe and assess the ankle for any
damage and to examine the collected data. The mode of injury, failure torque and failure
angle of foot rotation were documented for each specimen. Torsional stiffness was

26

determined based on the first 20° of external rotation [14]. Following the experimental
testing, a careful gross dissection of the joint was performed to ensure that all soft tissue
injuries and bone fractures were documented.

The average torque at various angles of rotation, and the average torsional stiffness were
used in subsequent statistical analyses. A Student’s t-test (unpaired, one-tailed) and a
Mann-Whitney test were separately conducted for comparison of torsional stiffness
between 0.5 Hz and 2.0 Hz frequencies of rotation, and between different angles of
flexion. The effects of rotation frequency and flexion angle on the torque-angle response
of the foot to external rotation were separately assessed by a one-way analysis of variance
(ANOVA) in SigmaStat (Version 2.03, SPSS Inc., Chicago, Illinois). SigmaStat was also
used to calculate the statistical β’s of the t-test and ANOVA. A Pearson product moment
correlation test was conducted to assess the correlations between the mode of injury and
rotational frequency, and between the mode of injury and flexion angles. Statistical
significance was set at p < 0.05 in all analyses.

RESULTS
The torsional stiffness for each frequency of rotation, 0.5 Hz (n=5) and 2.0 Hz (n=5), was
not significantly different (Figure 2-3, p > 0.05, β = 0.732). The mean torsional stiffness
for all specimens was 1.25 ± 0.46 Nm/deg (Table 2-1). The torque-angle response also
did not vary with frequency of foot rotation (p > 0.05, β =0.926). As for flexion angles of
20° (n=7), 10° (n=2), and -10° (n=1), no statistically significant differences in response

27

were noted (p > 0.05, β = 0.771 for torsional stiffness, and β = 0.805 for torque-angle

120

2

100

1.5
Torque (Nm)

Torsional Stiffness (Nm/deg)

behavior). The average torque-angle response for all specimens was shown in Figure 2-4.

1
0.5

80
60
40
20

0
0.5

0

2

0

Frequency of Rotation (Hz)

10

20

30

40

50

60

70

Angular Displacement (deg)

Figure 2-3 Torsional stiffness of foot/ankle

Figure 2-4 Average torque-angular

complex for different frequencies of

displacement response of foot/ankle

rotation. No significant differences between

complex for all specimens.

conditions were observed.

The mean failure torque of the ankles was 69.5 ± 11.7 Nm with a mean failure angle of
40.7 ± 7.3 degrees (Table 2-1). The various failure modes for different specimens were
listed in Table 2-1. Isolated avulsion of the PTaFL from the fibula was noted in four of
the ten ankles (Figure 2-5A-left). Distal oblique fibular fractures were generated in three
ankles, two of which passed through the anterior tibiofibular ligament (ATiFL) (Figure 25A-right). Single incidents were noted of midsubstance PTaFL rupture, anterior deltoid
ligament rupture (tibiotalar and tibionavicular) (Figure 2-5B) and a spiral fracture of the
tibia and fibula. The frequency of injury type was not a function of foot rotation
frequency (p = 0.073) or flexion angle (p = 0.204) in the current study. None of these

28

injuries were found to be associated with or adjacent to the screws that were inserted for
the motion study (data not reported here).

29

Table 2-1 – Torque and foot rotation at failure.
Rotation
Freq.
(Hz)

Flexion
(deg)

Torque
(Nm)

Angle
(deg)

0.5
2.0
2.0
0.5
0.5
0.5
2.0
2.0

20
20
20
20
20
20
10
10

55
60
67
95
61
62
68
76

34
35
31
55
45
35
40
45

86

0.5

20

80

47

64

2.0

-10

71

40

0.94

AVG
(SD)

69.5
(11.7)

40.7
(7.3)

1.25 (0.46)

Age

Ht
(m)

Wt
(kg)

70

1.52

86

34

1.83

68

47

1.47

113

40

1.78

113

32532L

19

1.88

32498R

47

1.78

Specimen
32388R
32388L
32416R
32416L
32516R
32516L
32489R
32489L

Torsional
Stiffness
(Nm/deg)
1.42
1.77
1.97
0.59
0.77
1.70
1.17
1.00
1.15

Injury Description
Fib avulsion of PTaFL
Fib avulsion of PTaFL
Distal fib fracture through ATiFL
Distal fib fracture
Distal fib fracture through ATiFL
Spiral fracture of tibia and fibula
Fib avulsion of PTaFL
Fib avulsion of PTaFL
Anterior deltoid (tibionavicular,
tibiotalar)
PTaFL midsubstance

PTaFL: posterior talofibular ligament. ATiFL: anterior tibiofibular ligament. Positive flexion is dorsiflexion.

30

A

Fibula

Fibula
ATiFL
Fibula fracture
through ATiFL
Fibula avulsion of PTaFL

B
Tibia

Tibiotalar
Anterior deltoid rupture

Figure 2-5 Representative failure modes resulting from excessive external rotation of the
foot/ankle complex. (A): Fibular avulsion of the posterior talofibular ligament (left);
fibular fracture through the anterior tibiofibular ligament (right). (B): Anterior deltoid
ligament rupture.

31

DISCUSSION
The new finding of this study was that during failure level laboratory experiments under
external rotation of the foot the most commonly injured structure was the PTaFL. This
finding was consistent with experimental studies that involve ligament sectioning and
prefailure strain measurements that document the relative importance of the PTaFL in
restraining external rotation of the foot [15,16]. Clinically, Colville et al. [16] also report
frequently finding tenderness in the PTaFL area in patients who have sustained
dorsiflexion-external rotation ankle injuries. In another clinical study over a thirty-three
month period Fallat et al. [18] document injury to the PTaFL in over thirty percent of
patients presenting with ankle sprains without osseous injury. Fallat et al. [18] also report
nine cases of isolated PTaFL injury. As indicated in a previous failure experiment by
Stiehl et al. [14], the ATiFL was only injured in the current study associated with a distal
oblique fibular fracture. In the current study the only injury to the anterior deltoid
ligament occurred in the ankle of the youngest specimen (19 years old). While additional
cadaver experiments may be needed, this result implied that age might influence the
location of ankle injuries.

The prevalence of posterior injuries in this series of experiments may be attributed to
constraint of the subtalar joint brought about by fixation of the calcaneous and the use of
potting material around the foot. A similar constraint of the foot/ankle complex was used
in the sectioning study that documents the important role of the PTaFL in the ankle’s
resistance to external rotation [15]. This constraint may affect the mechanics of the foot
and lead to excessive relative motion between the fibula and talus, and the resultant strain

32

in the PTaFL. Stiehl et al. [14] externally rotate the foot and document anterior deltoid
ligament injuries. The authors constrain the foot with fiberglass tape, thereby possibly
allowing subtalar motion. Another recent study examined the modes of injury when the
foot was constrained with athletic tape [21]. It was noted that the primary injury was to
the anterior deltoid ligament in three of four ankles. In the real world, this may imply that
shoes with tight, stiff uppers, as described by [20], may influence foot mechanics and
therefore the potential for changes in the location of soft tissue injury. Similar effects on
ankle injury, based on boot stiffness, have been previously documented in snowboarding
[22].

The mean failure torque in the current experiments, approximately 70 Nm, compared
well with the current literature. Hirsch and Lewis [13] document a mean failure torque of
75 ± 19 Nm to produce bone fractures (lateral and/or medial malleolus) in the ankle joint
during pronation-external rotation of the foot. Stiehl et al. [14] record a mean failure
torque for male specimens of 65 ± 21 Nm to generate oblique fibular fracture during
external rotation of the foot. In the current study the fibula related fractures were found in
four of ten specimens (Table 2-1). Additionally, Stiehl et al. [14] document an average
torsional stiffness of 1.21 ± 0.43 Nm/deg in male ankles under external rotation, while
the current study recorded an average torsional stiffness of 1.25 ± 0.46 Nm/deg.

Some limitations of the current study should be noted. Due to the relatively small sample
size used in the current study the statistical power (1-β) was low. Additional analyses
based on the current variances showed that with a power of 0.8 (β = 0.2) the sample size

33

would have to be more than 40 to be able to document significant differences in those
parameters. One unexpected limitation of the current study was the frequency of bone
avulsion injuries at the insertion of the PTaFL. Compared with other studies (average age
of the specimens was 64 and 67 years in Colville et al. [16] and Stiehl et al. [14],
respectively), we expected a younger aged specimen population (43

17 years) to exhibit

more soft tissue injuries. In fact, the higher frequency of foot rotation was employed to
try and enhance this possibility, but the mode of injury did not vary with frequency in the
current study. If the avulsion injuries were avoided, however, the anterior deltoid may
have become the “weakest link”. Other studies have shown, in fact, that the PTaFL is
stronger than portions of the anterior deltoid ligament [23,24]. If the bone quality would
have been more representative of the typical, young athlete, the anterior deltoid may have
been the location of primary failure. In fact, the single 19 years old specimen showed
gross injury in the anterior deltoid ligament in the current study. The number of avulsions
may still be attributed to the speed of testing, which was limited in the current study to
control for inertial effects of our current fixture. Stiehl et al. [14] externally rotated the
foot at a rate of 45° per second up to a magnitude of 90° and reported that 67% of male
specimens had an oblique fibular fracture, torn ATiFL, and torn anterior deltoid ligament.
Previous research on the knee suggests that the mode of ligamentous injury may be
dependent on the strain rate [25]. Avulsions are shown to occur more often at low rates of
loading, while midsubstance failures are frequently documented in high rate tests. In the
current study there were also not enough data in the 10° and -10° of flexion groups (n=2
and n=1, respectively), so documenting a flexion angle effect was limited. Another

34

limitation of the current study was that the foot restraint, in terms of the potential shoe
design issue, was not more completely studied at this time.

In conclusion, we found that during failure level, laboratory experiments, the PTaFL was
the most common injury site during external rotation of the foot with a relatively rigid
constraint. Visible damage to the syndesmosis only occurred in combination with fibular
fracture in these experiments. An anterior deltoid ligament injury was noted in the
youngest specimen. While clinically syndesmosis injury is often associated with external
foot rotation, the current study suggested that this fact might be due to a less constrained
subtalar joint in the real-world and a younger population diagnosed. In future
experiments the role of constraint offered to the foot and ankle complex by footwear with
varying upper rigidity should be investigated as it may affect the location and
torques/rotations needed to produce soft tissue injuries in the human ankle.

35

REFERENCES

36

REFERENCES

1. Boytim, M.J., Fischer, D.A., and Neumann, L., 1991, “Syndesmotic Ankle
Sprains,” American Journal of Sports Medicine, 19, pp. 294-298.
2. Guise, E.R., 1976, “Rotational Ligamentous Injuries to the Ankle in Football,”
American Journal of Sports Medicine, 4, pp. 1-6.
3. Johnson, E.E., and Markolf, K.L., 1983, “The Contribution of the Anterior
Talofibular Ligament to Ankle Laxity,” Journal of Bone and Joint Surgery
(American Volume), 65, pp. 81-88.
4. Rasmussen, O., Tovborg-Jensen, I., and Boe, S., 1982, “Distal Tibiofibular
Ligaments,” Acta Orthopaedica Scandinavica, 53, pp. 681-686.
5. Xenos, J.S., Hopkinson, W.J., Mulligan, M.E., Olson, E.J., and Popovic, N.A.,
1995, “The Tibiofibular Syndesmosis: Evaluation of the Ligamentous Structures,
Methods of Fixation, and Radiographic Assessment,” Journal of Bone and Joint
Surgery (American Volume), 77, pp. 847-856.
6. Bahr, R., Pena, F., Shine, J., Lew, W.D., and Engebretsen, L., 1998, “Ligament
Force and Joint Motion in the Intact Ankle: A Cadaveric Study,” Knee Surgery,
Sports Traumatology, Arthroscopy, 6, pp. 115-121.
7. Shybut, G.T., Hayes, W., and White, A.A., 1983, “Normal Pattern of Ligament
Loading Among the Lateral Collateral Ankle Ligaments,” Transactions of the
Orthopaedic Research Society, 8, pp. 15.
8. Edwards, G.S., and DeLee, J.C., 1984, “Ankle Diastasis without Fracture,” Foot
and Ankle, 4, pp. 305-312.
9. Hopkinson, W.J., St. Pierre, P., Ryan, J.B., and Wheeler, J.H., 1990,
“Syndesmosis Sprains of the Ankle,” Foot and Ankle, 10, pp. 325-330.
10. Lauge-Hansen, N., 1950, “Fractures of the Ankle II. Combined ExperimentalSurgical and Experimental-Roentgenologic Investigations,” Archives of Surgery,
60, pp. 957-985.
11. Markolf, K.L., Schmalzried, T.P., and Ferkel, R.D., 1989, “Torsional Strength of
the Ankle in Vitro. The Supination-External-Rotation Injury,” Clinical
Orthopaedics and Related Research, 246, pp. 266-272.
12. Schaffer, J.J., and Manoli, A., 1987, “The Antiglide Plate for Distal Fibular
Fixation,” Journal of Bone and Joint Surgery (American Volume), 69, pp. 596604.

37

13. Hirsch, C., and Lewis, J., 1965, “Experimental Ankle Joint Fractures,” Acta
Orthopaedica Scandinavica, 36, pp. 408-417.
14. Stiehl, J.B., Skrade, D.A., and Johnson, R.P., 1992, “Experimentally Produced
Ankle Fractures in Autopsy Specimens,” Clinical Orthopaedics and Related
Research, 285, pp. 244-249.
15. Stormont, D.M., Morrey, B.F., An, K.N., and Cass, J.R., 1985, “Stability of the
Loaded Ankle. Relation between Articular Restraint and Primary and Secondary
Static Restraints,” American Journal of Sports Medicine, 13, pp. 295-300.
16. Colville, M.R., Marder, R.A., Boyle, J.J., and Zarins, B., 1990, “Strain
Measurements in Lateral Ankle Ligaments,” American Journal of Sports
Medicine, 18, pp. 196-200.
17. Rasmussen, O., 1985, “Stability of the Ankle Joint. Analysis of the Function and
Traumatology of the Ankle Ligaments,” Acta Orthopaedica Scandinavica, 56
(Suppl 211), pp. 1-75.
18. Fallat, L., Grimm, D.J., and Saracco, J.A., 1998, “Sprained Ankle Syndrome:
Prevalence and Analysis of 639 Acute Injuries,” Journal of Foot and Ankle
Surgery, 37, pp. 280-285.
19. Molus, M.A., and Martin, D.F., 2008, “Talofibular Ligament Injury,” Retrieved
December
17,
2008,
from
eMedicine
WebMD
Web
site:
http://emedicine.medscape.com/article/86396-overview.
20. Villwock, M.R., Meyer, E.G., Powell, J.P., and Haut, R.C., 2009a, “Football
Playing Surface and Shoe Design Affect Rotational Traction,” American Journal
of Sports Medicine, 37, pp. 518-525.
21. Villwock, M.R., Meyer, E.G., Powell, J.P., and Haut, R.C., 2009b, “External
Rotation Ankle Injuries: Investigating Ligamentous Rupture [abstract],” In:
Proceedings of 2009 ASME Summer Bioengineering Conference; Jun 17-21;
Lake Tahoe, CA.
22. Delorme, S., Tavoularis, S., and Lamontagne, M., 2005, “Kinematics of the Ankle
Joint Complex in Snowboarding,” Journal of Applied Biomechanics, 21, pp. 394403.
23. Butler, A.M., and Walsh, W.R., 2004, “Mechanical Response of Ankle Ligaments
at Low Loads,” Foot and Ankle International, 25, pp. 8-12.
24. Siegler, S., Block, J., and Schneck, C.D., 1988, “The Mechanical Characteristics
of the Collateral Ligaments of the Human Ankle Joint,” Foot and Ankle, 8, pp.
234-242.

38

25. Crowninshield, R.D., and Pope, M.H., 1976, “The Strength and Failure
Characteristics of the Rat Medial Collateral Ligament,” Journal of Trauma, 16, pp.
99-105.

39

CHAPTER 3
DEVELOPMENT AND VALIDATION OF A COMPUTATIONAL MODEL TO
STUDY THE EFFECT OF FOOT CONSTRAINT ON ANKLE INJURY DUE TO
EXTERNAL ROTATION

ABSTRACT
Recent studies, using two different manners of foot constraint, potted and taped,
document altered failure characteristics in the human cadaver ankle under controlled
external rotation of the foot. The posterior talofibular ligament (PTaFL) was commonly
injured when the foot was constrained in potting material, while the frequency of deltoid
ligament injury was higher for the taped foot. In the current study an existing multibody
computational modeling approach was validated to include the influence of foot
constraint, determine the kinematics of the joint under external foot rotation, and
consequently obtain strains in various ligaments. It was hypothesized that the location of
ankle injury due to excessive levels of external foot rotation is a function of foot
constraint. The results from this model simulation supported this hypothesis and helped
to explain the mechanisms of injury in the cadaver experiments. An excessive external
foot rotation might generate a PTaFL injury for a rigid foot constraint, and an anterior
deltoid ligament injury for a pliant foot constraint. The computational models may be
further developed and modified to simulate the human response for different shoe designs,
as well as on various athletic shoe-surface interfaces, so as to provide a computational
basis for optimizing athletic performance with minimal injury risk.
Keywords: ankle sprain, ligament, strain, three-dimensional reconstruction, simulation,
subtalar joint, arch stability, joint mechanics, motion analysis, joint coordinate system.

40

INTRODUCTION
Mechanisms of ankle injury to the syndesmosis3,14,27 and lateral ligaments6,19,26 have been
investigated in numerous studies. In the National Football League (NFL), syndesmotic
sprains constitute approximately 20% of ankle injuries3. Syndesmotic sprains are
particularly problematic because of the long period needed for rehabilitation, as they
involve injury to any or all of the following soft tissue structures: the anterior tibiofibular
ligament (ATiFL), the posterior tibiofibular ligament (PTiFL) and the interosseous
ligament27,33. Lateral ankle injuries are related to the ATiFL, the PTiFL, the
calcaneofibular (CaFL), the anterior talofibular (ATaFL) and the posterior talofibular
(PTaFL) ligaments6. The deltoid ligament is also frequently injured in combination with a
syndesmotic sprain.7,8,24,45 The mechanisms associated with this injury may involve foot
eversion,9,39,40 dorsiflexion,4,14,28 and external rotation.3,27,45

Sectioning studies have described the role of ligaments to resistive torque during external
foot rotation.16,30,31,38,46 This involves sequentially cutting various ligaments and
quantifying their individual contribution to joint stiffness. Buckle transducers and strain
gauges have also been used to show the contributions of various ligaments to external
rotation of the foot.1,6,34 A potential problem with ligament sectioning-type studies is that
ankle biomechanics may be altered and the findings may not adequately describe
dynamic relationships between ligaments.6 And, studies with gauges on ligaments
generally are not performed at ligament failure strain levels.

41

Research by Lundberg et al.21,22 suggests that rotations of the bones within the foot
during external rotation may act as torsion dissipating devices. This might imply that
injury mechanisms may be influenced by the constraint of various bones in the foot. Two
methods of foot constraint have been used to study the biomechanical response and
tolerance of the ankle to external foot rotation. Stormont et al.38 rigidly fix the foot in a
potting alloy and perform sectioning experiments. The study concludes that the PTaFL
and CaFL are the primary restraints to external foot rotation. In contrast, Stiehl et al. 37
constrain the foot with fiberglass cast tape. The study produces injury to the anterior
ligamentous structures, including the deltoid ligament, ATiFL, and interosseous ligament
in association with fracture of the fibula and tibia. These contrasting data on locations of
primary restraint during external foot rotation may be the result of differences in foot
constraint. Foot constraint may influence subtalar motion and the movement of bones in
the foot, thereby influencing the mode of injury during external foot rotation. Recent
studies41,44 in our laboratories, using potted and taped foot constraints, also document
altered failure characteristics of the ankle under external foot rotation. In a study with a
potted foot, the PTaFL was injured in 5 of 10 ankles; fibular fractures were generated in 4
ankles; a single incident of an anterior deltoid ligament rupture was noted.44 In a study
with taped foot constraint, the deltoid ligament was damaged in 3 of 4 ankles; fibular
fracture was noted in 1 ankle.41

Computational models of musculoskeletal joints and limbs can provide useful
information about joint mechanics.20 Validated models can be predictive tools to
understand normal joint function and serve as clinical tools for predicting and helping to

42

prevent sports injuries. While finite element modeling yields useful information about
stresses and strains in bones and ligaments,5,32 an advantage of multibody rigid modeling
is the ability to rapidly solve for motion-based mechanics in large structures.17 Liacouras
and Wayne20 have developed a computational approach to model the lower leg to
simulate cadaver studies of syndesmotic injury and ankle inversion stability. A similar
model has been validated with an application to arch stability by Iaquinto and Wayne. 15
While both studies demonstrate an encouraging advance of this modeling approach, they
only simulate ligament sectioning experiments, and motions of various joints are not fully
controlled. By using this existing computational approach, a graphic-based, multibody
computational model of the foot/ankle complex was developed and validated in the
current study to investigate joint motion, and therefore to study the effect of foot
constraint on the response of the ankle to external rotation. In addition, the previous
cadaver experiments41,44 were revisited to obtain talus rotations for comparison with the
model. It was hypothesized that the mechanisms and specific locations of ankle injury in
previous cadaver experiments, using potted and taped foot constraints, could be simulated
with a multibody computational model using parameter information from the current
literature.

43

METHODS
Bone reconstruction
For the current study a fresh frozen, left cadaver foot (19 years old) showing no signs of
abnormal anatomy, transected approximately 15 cm distal to the knee center, was fully
thawed overnight and CT scanned to obtain three-dimensional joint anatomy. Detailed
features of the ankle were transferred from Digital Imaging and Communications in
Medicine (DICOM) files into Materialise’s Interactive Medical Imaging Control System
(MIMICS) (Materialise, Ann Arbor, MI). The individual bones of the foot were
computationally separated and meshed as solid bodies. This yielded a 3-D surface model
of each bone as Stereolithography (STL) files for export. To reduce the size of surface
files and subsequent models, STL files were remeshed in MIMICS to smooth the surface
of each bone. The remeshed STL files were then imported into a 3-D solid modeling
program (SolidWorks, TriMech Solutions, LLC, Columbia, MD). In SolidWorks, the
ScanTo3D package was used to reconstruct each bone and simplify the bone surfaces.
The individual bones were then assembled into a lower extremity and registered with
respect to one another in the anatomical position (Figure 3-1). Since toe involvement is
minimal in the experimental cadaver foot, the phalanges were excluded in the model. All
other joints of the foot and ankle were constrained only by joint geometry and ligament
stability.

44

B

A

Figure 3-1 Posterior (A) and lateral (B) views of a reconstructed 3D model of the bony
foot anatomy. Model was scanned and reconstructed in a neutral anatomical orientation
(approximately 90° between the foot plane and the lower leg). The bones of the foot were
connected by linear springs as ligaments. The phalanges (toes) were excluded in the
model. For interpretation of the references to color in this and all other figures, the reader
is referred to the electronic version of this dissertation.

45

Table 3-1 Ligament abbreviations and stiffness values used for each spring (ligament)
present in the models.
Ligaments represented in models

Abbreviation

Stiffness (N/mm) [Ref]

Interosseous I

Inte-I

400 [29]

Interosseous II

Inte-II

400 [29]

Anterior Tibiofibular

ATiFL

90 [35]

Posterior Tibiofibular

PTiFL

90 [2]

Calcaneofibular

CaFL

70 [20]

Anterior Talofibular

ATaFL

90 [35]

Posterior Talofibular

PTaFL

70 [20]

Anterior Tibiotalar

ATiTL (Deltoid)

70 [20]

Posterior Tibiotalar

PTiTL

80 [20]

Tibionavicular

TiNL

40 [35]

Talonavicular

TaNL

70 [20]

Interosseous Talocalcaneal

ITaCL

70 [20]

Lateral Talocalcaneal

LTaCL

70 [20]

Medial Talocalcaneal

MTaCL

70 [20]

Posterior Talocalcaneal

PTaCL

70 [20]

Calcaneocuboid

CaCL

70 [20]

Calcaneonavicular

CaNL

70 [20]

Medial Plantar Fascia

MPF

200 [15]

Central Plantar Fascia

CPF

230 [15]

Lateral Plantar Fascia

LPF

180 [15]

46

Model formulation
SolidWorks Motion (SolidWorks, TriMech Solutions, LLC, Columbia, MD) was
implemented to apply ligamentous restraints, prescribe force/motion constraint, and
simulate the ankle dynamics. Ligaments were represented as linear spring elements
(Figure 3-1), with stiffness values from the literature (Table 3-1). Origin and insertion
locations of the ligaments were determined from dissection and anatomical atlases.23,25
Ligament preloads were induced by reducing lengths by 2%.20 Each bone was allowed to
move in all six degrees of freedom, leaving body motion to be a function of ligament
behavior, surface contact, and external constraints. Similar to the methodology used by
Iaquinto and Wayne,15 the 3-D articular contact conditions were simulated between
adjacent bone surfaces in COSMOSMotion of SolidWorks. Using the centroid of each
interfering volume as the point of contact, equal and opposite reaction forces were
applied to each body along the contact normal direction. Friction and the effects of
gravity were considered negligible. The tibia was fixed in space for all simulations.
Simulated body load was applied to the proximal end of the model, proportionally
distributing it between the tibia and the fibula as to one-sixth loading on the fibula.18,43

Model validation – ligament strain
The first validation recreated the experimental cadaver study performed by Colville et al.6
to investigate the strain present in lateral ligaments of human ankles while moving the
ankle joint and applying stress in a variety of ways. As done experimentally, the ankles
were moved from 20° of dorsiflexion to 30° of plantar flexion while no
inversion/eversion or rotational forces were applied. Strain was measured continuously,

47

by Colville et al.,6 in each ligament using strain gauges through this range of motion.
Strain was also measured continuously in each ligament through this range of motion
while 3 Nm of internal and external torque was applied to the ankle. Subtalar joint motion
was eliminated in this experimental study by the use of two screws passing through the
calcaneus into the talus. To mimic this experimental foot constraint condition, the tibia in
the computational model was fixed while the fibula was free to move, and the rest of the
bones were fused together and moved as a rigid body (Figure 3-2A). Ten ligaments were
used to connect bones in this model; namely, the Inte-I, Inte-II, ATiFL, PTiFL, CaFL,
ATaFL, PTaFL, ATiTL, PTiTL, and TiNL (Table 3-1). To simulate this cadaver study
with the computational model, continuous dorsi-plantar flexion was applied to the foot
with the axis of rotation through the talus. A rotational torque of 3 Nm was then
internally or externally applied to the talus with the axis of rotation along the tibia for
various dorsi-plantar flexion angles. Ligament strains in the ATiFL and PTiFL were
calculated in the model and compared to the cadaver data. No compressive load was
applied in this computational model, as done experimentally.

48

A

B

C

Figure 3-2 Various models used in the validation studies and the foot constraint study. In
all models, the tibia was fixed in space, and the bones in blue color were fused together
and moved as a rigid body. Model (A) represented a fused subtalar joint for the ligament
strain validation. Model (B) allowed the subtalar joint motion, representing the potted
foot in the external rotation validation. Model (C) further allowed the arch collapse,
representing the taped foot in the foot constraint study.

Model validation – external foot rotation
A second validation of the ankle model was performed while simulating a cadaver study
by Wei et al.44 that investigates ankle injury due to excessive external foot rotation. This
cadaver study used a relatively rigid foot constraint by placing screws into the calcaneus
and by potting the foot in epoxy resin. While this foot constraint prevented the arch from
collapsing, subtalar joint motion was allowed. The foot was positioned in approximately
20° of dorsiflexion and 10° of eversion in these experiments. In addition, a compressive
preload of 2000 N was applied to the ankle through the proximal end of the tibia and
fibula. External rotation was increased by increments of 5° in successive tests until
diagnosed ligamentous injury or bone fracture. The injury mode, failure torque, failure
49

foot rotation, and torsional stiffness were documented in the study. To simulate this foot
constraint, the talus was set free in the computational model to allow subtalar joint
motion (Figure 3-2B). Consequently, 5 more ligaments were incorporated in the model;
namely, the TaNL, ITaCL, LTaCL, MTaCL, and PTaCL (Table 3-1). To simulate this
cadaver experiment in the computational model, an input external foot rotation of 40°
was used, based on the average failure rotation documented in the cadaver study.
Compressive loads of 1660 N and 340 N were respectively applied on the tibia and the
fibula in the model, assuming approximately one-sixth of loading carried by the
fibula.18,43

An additional variable that controls the kinematics of the ankle joint is the stiffnesses of
ligaments. Reported ligament stiffness values have a relatively large range.2,15,20,29,35 To
assess the sensitivity of the simulations and further validate the computational models,
the effect of variations in ligament stiffness on model outcomes was studied by
performing an additional simulation that mimicked the experimental setup of Wei et al.44
To quantify the effect of ligament stiffness, first, all ligament spring stiffnesses were
together either increased by 25% or decreased by 25%. Second, strains of each ligament
at 40° external foot rotation were compared and those ligaments that had more than 10%
strain were selected to individually vary their stiffnesses by +/-25%. Finally, the selected
ligaments were randomly combined and variation of stiffness by +/-25% was performed
at the same time.

50

Foot constraint study and injury prediction
The model from the second validation study was used to simulate the cadaver study44
involving a potted foot constraint. In contrast, Villwock et al.,41 using the same
experimental protocol as the Wei et al.44 study, fixed the foot with athletic tape and
documented primarily injuries to the anterior deltoid ligament. The results are consistent
with a previous study by Stiehl et al.37 To simulate this relatively less rigid foot constraint,
a taped foot constraint, the current model was modified by releasing the calcaneus to
allow arch collapse, while including elements representing the plantar fascia to maintain
arch stability (Figure 3-2C). Consequently, 5 more ligaments were added, namely, the
CaCL, CaNL, MPF, CPF, and LPF (Table 3-1). An external foot rotation of 65° was
applied as input in the taped foot computational model, based on the average failure
rotation documented by Villwock et al.41 Again, compressive loads of 1660 N and 340 N
were applied to the tibia and the fibula, respectively. Ligament strains and talus rotation
relative to the tibia were calculated in the computational models.

In the cadaver studies by Wei et al.44 and Villwock et al.41, four specimens from each
group were subjected to talus motion analysis (data not presented previously) using a sixcamera Vicon motion capture system (Oxford Metrics Ltd., Oxford, United Kingdom).
To compare the talus rotations in the computational simulations with cadaver studies, the
previous Vicon data were re-analyzed for the current study. Grood and Suntay11
suggested forming a joint coordinate system (JCS) using Euler angles for clinical
description of 3-D motion of the knee. This was later applied to the ankle by Soutas-Little
et al.36 to measure 3-D rotation of the foot relative to the shank. The JCS36 required

51

establishing two local coordinate systems; in their case, one attached to the foot and one
to the shank. The JCS took one coordinate axis from each of these two local systems and
the final floating coordinate axis or line of nodes as perpendicular to the other two. The
JCS is not, in general, an orthogonal coordinate system but this does not present
analytical problems.11,36 In the current study a JCS was established based on a reflective
marker triad fixed to the talus on the anterior aspect of the cadaver foot (Figure 3-3). The
two local coordinate systems were chosen as the same marker triad before and after the
external foot rotation. Clinical angles of talus motion relative to the tibia were calculated
at 35° of foot rotation in the ankle JCS. This analysis was performed in each experiment
preceding gross joint failure. These angles could be compared directly to the talus
rotations from the models.

Figure 3-3 Schematic of the cadaver foot with a marker set screwed into the talus on the
anterior aspect of the foot.

52

20
No torque (Colville)

A

External torque (Colville)

ATiFL

Internal torque (Colville)
No torque (Model)

Ligament Strain (%)

15

External torque (Model)
Internal torque (Model)

10

5

0
-20

-10

0

10

20

30

20
No torque (Colville)

B

External torque (Colville)

PTiFL

Internal torque (Colville)
No torque (Model)

Ligament Strain (%)

15

External torque (Model)
Internal torque (Model)

10

5

0
-20

-10

0

10

Dorsiflexion (-)

20

Plantarflexion (+)

30
(deg)

Figure 3-4 Comparison between computational model and cadaver experiment6 of
ligament strains in the ATiFL (A) and the PTiFL (B).
53

RESULTS
Ligament strain
Ligament strains at various dorsi-plantar flexion angles were documented experimentally,
either with or without an applied torque.6 In the computational model, strain in the ATiFL
increased 2.5% when the ankle was moved from the neutral position to 20° of
dorsiflexion (Figure 3-4A). While external torque on the talus increased strain by
approximately 2%, internal torque on the talus decreased strain. This was true for all joint
angles. The PTiFL elongated up to a maximum of 2% as the ankle was moved from the
neutral position to 20° of dorsiflexion (Figure 3-4B). Strain in the ligament decreased
with external torque and increased with internal torque on the talus. Ligament strains in
the model during this range of motion testing followed trends seen experimentally.6

80
Cadaver study
Computer model

Torque (Nm)

60

40

20

0
-10

0

10

20

30

40

50

Rotation (deg)
Figure 3-5 Comparison of torque-rotation behavior from computational model to cadaver
experiment.44 The curve from the model was shifted to the left to account for 8.5° of
internal rotation generated by a compressive load applied to the model in a pilot study.

54

External foot rotation
Comparison of the computational model to the cadaver study44 demonstrated a similar
torque-rotation behavior during external rotation of the foot (Figure 3-5). In a preliminary
model, an axial load of 2000 N on the ankle model resulted in the foot internally rotating
8.5°. The torque-rotation curve for the model was therefore shifted 8.5° to account for
this internal rotation. While seen experimentally, this pre-load effect was not documented
in the previous study. The cadaver study also showed an average failure torque of 69.5
Nm with a standard deviation (SD) of 11.7 Nm, an average failure rotation of 40.7° with
a SD of 7.3°, and an average torsional stiffness of 1.25 Nm/deg with a SD of 0.46
Nm/deg. The computational model showed a similar maximum torque of 60.1 Nm at 40°
of rotation (Figure 3-6A), and a torsional stiffness of 1.29 Nm/deg (Figure 3-6B) for the
average ligament stiffnesses used in the literature.
100

2

80

A

Rotational Stiffness (Nm/deg)

Maximum Torque (Nm)

-25%
Avg
+25%

60

40

20

-25%

B

Avg

1.6

+25%

1.2

0.8

0.4

0

0
Cadaver

Cadaver

Model

Model

Figure 3-6 Maximum torque (A) and rotational stiffness (B) induced by external foot
rotation from cadaver study44 and computational model simulation. Model sensitivity
tests were conducted where ligament stiffness values were 25% below and above the
average for selected ligaments.
55

At 40° of external foot rotation, the PTaFL, ATiTL and TiNL generated more than 10%
strains (Figure 3-7). Therefore, these three ligaments were selected to perform the
sensitivity tests. Sensitivity tests, parametrically varying the stiffnesses of the ankle
ligaments, resulted in small changes in maximum torque and rotational stiffness (Table 32). A 25% decrease in ligament stiffnesses yielded a 2.2% decrease in torque at 40°
rotation, and an 8.5% decrease of the ankle torsional stiffness. Conversely, raising
ligament stiffnesses by 25% produced a 3.8% increase in torque, and a 5.4% increase in
ankle torsional stiffness. The means and standard deviations of torque and torsional
stiffness across the ligament combinations (Table 3-2) were plotted in Figure 3-6 for
comparison with the experimental results. While these results demonstrated a trend for a
stiffer joint with stiffer ligaments, both parameter variations caused a change lower than
the experimental standard deviations reported by Wei et al.44 This suggested that
moderate parameter variations do not significantly alter conclusions that could be drawn
from this model using existing data from the literature.
40
35

Strain (%)

30
25
20
15
10
5

C
L

C
L

L

aC
L
PT

M
Ta

LT
a

IT
aC

Ta
N
L

Ti
N
L

iT
L
PT

Ti
TL
A

aF
L
PT

Ta
FL
A

aF
L
C

iF
L
PT

-II

Ti
FL
A

In
te

In
te

-I

0

Ligaments

Figure 3-7 Strains in different ligaments of the potted foot model at 40° of external
rotation. Ligaments with more than 10% strain were selected for the sensitivity analysis.
56

Table 3-2 Sensitivity analysis results. Three ligaments (PTaFL, ATiTL, and TiNL) were selected based on ligament strains at 40°
external foot rotation shown in Figure 3-7.
Ligament Combinations
PTaFL
TiNL

ATiTL
TiNL

PTaFL
ATiTL
TiNL

All
Ligaments

Mean

SD

PTaFL

Torque
(Nm)
Torsional
Stiffness
(Nm/deg)

ATiTL

TiNL

PTaFL
ATiTL

-25%

59

59

60

58

59

59

58

58

58.75

0.71

+25%

61

60

60

64

63

61

65

65.2

62.40

2.17

-25%

1.2

1.2

1.3

1.1

1.2

1.2

1.1

1.1

1.18

0.07

+25%

1.4

1.3

1.3

1.4

1.4

1.3

1.4

1.4

1.36

0.05

57

Foot constraint study and motion analysis
Ligament strains were determined for the potted and taped computational models. For
clarity, only strains in the 3 ligaments injured in the cadaver experiments were plotted.
While linear springs were used, the ligament strain-rotation behavior was non-linear
(Figure 3-8). All 3 ligaments experienced higher strains for the potted than the taped foot
model at the same foot rotation. At each level of rotation the computational model for the
potted foot showed the largest strain in the PTaFL (Figure 3-8A), while the deltoid
ligament generated the largest strain for the taped foot (Figure 3-8B). These results
paralleled with the cadaver studies, where the PTaFL44 and the deltoid ligament,41
respectively, were the tissues that actually failed experimentally for each type of
constraint. The ATiFL experienced the lowest strain for both models, and did not show
any gross injuries experimentally.
60

A
Ligament Strain (%)

60

Potted Foot Model

Taped Foot Model

B

50

50
PTaFL
Deltoid
ATiFL

40

PTaFL
Deltoid
ATiFL

40

30

30

20

20

10

10

0

0
0

10

20

30

40

50

60

70

External Rotation (deg)

0

10

20

30

40

50

60

70

External Rotation (deg)

Figure 3-8 Strains in different ligaments during external rotation of the computational
foot models: (A) potted foot; (B) taped foot. For clarity, only strains of the ligaments
injured in the cadaver experiments were plotted.
58

Based on the JCS established for re-analysis of cadaver data, the 3-D rotations of the
talus relative to the tibia were described in clinical terms as dorsi/plantar flexion,
eversion/inversion, and external/internal rotation. In these cadaver tests 35° of potted foot
rotation generated 33.7° ± 0.5° of talus rotation, and 35° of taped foot rotation resulted in
32.7° ± 1.8° of talus rotation experimentally (Figure 3-9). These results indicated the
accuracy of the JCS used in the analysis. External rotation also resulted in 4.7° ± 2.5° of
talus dorsiflexion and 0.5° ± 0.8° of talus inversion for the potted foot, and 0.7° ± 3.1° of
talus plantarflexion and 2.5° ± 4.1° of talus eversion for the taped foot experimentally.
Comparably, in the computational models 35° of external rotation for the potted foot
model generated 14º of talus dorsiflexion and 3º of talus inversion, while 35° of external
rotation for the taped foot model induced 5º of talus plantarflexion and 8º of talus
eversion relative to the tibia (Figure 3-9).

59

Talus Rotation Relative to the Tibia
(deg)

40

Dorsi(+) /
Plantar(-) Flexion

Eversion(+) /
Inversion(-)

External Rotation

30
Cadaver Study
Computer Model

20

10

0
Potted

Taped

Potted

Taped

Potted

Taped

-10
Foot Constraint

Figure 3-9 Talus rotations relative to the tibia obtained from the cadaver studies and the
computational models at 35° of external rotation for both the potted and the taped foot.

DISCUSSION
In the current study a computational foot model was developed and validated against two
well-documented cadaver studies. The validated model was then used to explore the
effect of two different types of foot constraint on ankle ligament strains and talus motions.
By allowing subtalar joint motion and arch collapse, the computational models showed
their ability to simulate different foot constraints used experimentally.37,38,41,44 The results
from this model analysis supported the hypothesis that foot constraint may affect ankle
behavior under external rotation, and lead to different locations of potential ankle injury.
The foot constraint study here paralleled some literature findings. Colville et al. 6 used
mercury-filled Silastic strain gauges to measure strain in lateral ligaments of 10 human
60

cadaver ankles with a relatively rigid foot constraint. The study shows extreme
dorsiflexion and external rotation of the talus increases PTaFL strain. This is consistent
with their clinical experience of frequent tenderness in the PTaFL area of patients with
dorsiflexion-external rotation ankle injury. In the current study the talus experienced
dorsiflexion and inversion for the potted foot computational model (Figure 3-9),
stretching the posterior-lateral aspect of the ankle and generating high strain in the PTaFL
(Figure 3-8A). In contrast, the high rate of deltoid injury in Villwock et al.41 for the taped
foot may have been due to subtalar motion that caused the arch to flatten resulting in talus
eversion and plantarflexion (Figure 3-9), stretching the anterior-medial aspect of the
ankle and generating high strain in the anterior deltoid ligament (Figure 3-8B). And,
Stormont et al.38 conducted serial sectioning tests and has shown that 83% of ankle
resistance to eversion is provided by the deltoid ligament.

Some studies have shown that the PTaFL is stronger than portions of the anterior deltoid
ligament.10,35 Funk et al.10 document the failure load of the PTaFL is 554.2 ± 94.6 N.
However, little has been reported about the strength of the individual structures which
make up the deltoid ligament, as most authors have not distinguished them from each
other. A study by Beumer et al.2 determine the strength of the tibiofibular and tibiotalar
ligaments of the ankle and show no difference between these ligaments having an
average strength of 550 ± 82 N. However, Funk et al.10 document the failure load of the
ATiTL is 130.8 ± 2.0 N. To quantify the influence of a stiffer PTaFL on its strain,
additional simulations were conducted in the current study and the results showed that

61

individually increasing the PTaFL stiffness by 50% did not change the strain patterns
observed in Figure 3-8.

Some limitations and assumptions of the model should be noted. Linear spring elements
were used as ligaments in the current model. Since ligament behavior is known to be
nonlinear, when computational models are used for failure analysis in the future,
nonlinear spring representation of ligaments should be considered for these large strains.
Although results in Figure 3-9 indicated that the talus rotations in the computational
model followed trends observed in the cadaver study, they were 3-5 times greater in
magnitude (except for the “external rotation" group). In the model reconstruction,
individual bones were smoothed, leaving bone motion to be dependent only on ligament
behavior and surface contact. This may allow more motion of individual bones in the
model than in reality during foot rotation and lead to the differences in Figure 3-9. A
model sensitivity test was conducted in the current study by (1) individually varying the
stiffnesses of the 3 highest strained ligaments (PTaFL, ATiTL, and TiNL in Figure 3-7)
and (2) varying their stiffnesses randomly in various combinations (Table 3-2). While a
sensitivity test with all 15 ligaments involved would lead to numerous possible
combinations and might be infeasible, a test with 4 ligaments (the above 3 and ATiFL)
was conducted and no significant difference from the results in Table 3-2 was observed.
And, the torque and torsional stiffness for the addition of the lower strained ligament did
not exceed the variance in the experimental data (Figure 3-6). While in experiments rigid
constraint at the proximal end of the tibia and fibula might have limited relative motion
between these two bones, the bones were likely deformed during the experiments, as

62

evidenced by previous studies producing bone fractures.37,41,44 While the bones were
assumed to be rigid and undeformable in these computational models, the fibula was
freed from the tibia to allow relative motion. Compared to the magnitudes of bone motion
and ligament elongation in the current model, the deformation of articular cartilage was
negligible. Cartilage function was incorporated by neglecting friction in the model. 20 And
finally, while the anterior deltoid ligament is actually composed of the ATiTL and
TiNL,25 in the current models the deltoid ligament was represented by only the ATiTL
(Table 3-1). The effect on more precisely discriminating localized injuries in the deltoid
complex itself is currently unknown.

Subtalar motion and arch collapse were allowed in the taped foot model by freeing up the
calcaneus (Figure 3-2C). Consequently, the ligaments in the taped foot experienced less
strain than in the potted foot for the same degree of rotation (Figure 3-8). This might
suggest that a more constrained foot might generate ankle ligament injury at a lower foot
rotation. A shoe with a pliant upper may allow more subtalar motion and natural
movement of bones in the foot so as to help protect it from injury. For performance
purposes, however, football and soccer players tend to wear shoes that provide “high
traction and good stability for fast accelerations, stops and turns”.13 Based on the current
model analysis, this might compromise the ankle mechanics and increase the risk of
injury. High rotational tractions between shoes and the playing surface may also
contribute to an increased frequency of ankle injury.12 Villwock et al.,42 for example,
used a mobile testing apparatus to measure shoe-surface interface torques for various
cleat patterns and playing surfaces. Artificial surfaces yielded higher torques and

63

rotational stiffnesses of the ankle than on natural grass surfaces. The authors42 suggest
rotational stiffness may largely depend on rigidity of the shoe’s upper. Developing
computational models of the shoe-surface interface conditions and the upper of shoes, for
example, may provide a more cost-effective means of trying to optimize player
performance while minimizing injury risk.

In conclusion, the ability of the current models to predict the behavior of cadaver
experiments was encouraging. The hypothesis of this study was supported using these
models showing that foot constraint may influence the potential for and determine the
location of soft tissue, ankle injury during external foot rotation. By changing the
magnitude of subtalar joint motion so as to match foot constraint offered by shoes, the
computational models may be further developed and modified in the future to simulate
the human response for various shoe designs and on different athletic shoe-surface
interfaces. These models may provide a useful computational tool for optimization of
shoe and shoe-surface interfaces in the future.

64

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65

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69

CHAPTER 4
SIMULATION STUDIES TO INVESTIGATE ANKLE SPRAIN MECHANISMS
PART I: MECHANISM OF INJURY IN A LATERAL ANKLE SPRAIN

INTRODUCTION:
Ankle sprain is the most common injury in sports [3], but the mechanism of injury may
not always be clear. Sports-related injury mechanisms can be studied through various
approaches: cadaveric experiments, kinematics, biomechanics, and computational
modeling. The advantage of computer simulation is the ability to separate motions and
study them individually.

In a recent study by Wei et al. [7] a computational foot model has been developed based
on computed tomography (CT) scans and validated against two well-documented cadaver
studies [2,8]. The validated model has then been used to explore the effect of two
different types of foot constraint, rigid and compliant, on ankle ligament strains and talus
motions. The model has shown its ability to predict ankle injuries due to excessive levels
of external foot rotation.

Recently, Fong et al. [3] reported a case study where an accidental supination ankle
sprain injury occurred in a laboratory under a high-speed video and plantar pressure
capturing setting. The injury was diagnosed as a grade I, mild anterior talofibular
ligamentous (ATaFL) sprain. Calcaneofibular ligament (CaFL) and syndesmotic
involvement were ruled out. The ankle kinematics of the injury trial documents a large
amount of ankle inversion, and an analysis of the injury video sequences shows some

70

ankle plantar flexion. These provide information for understanding the ankle sprain
mechanism quantitatively. Interestingly, the authors also describe a pivoting internal
rotational motion at the moment of injury, where the forefoot was in touch with the
ground and supported the body, but the rearfoot drifted to the lateral side. This implies
that internal rotation and inversion may both have been involved in the ankle sprain. The
purpose of the current study was to test the influence of various motions on ankle
ligament strains using a multibody three-dimensional simulation model driven by in vivo
data from the injury event. The hypothesis was that internal rotation and plantar flexion
applied to the ankle, in addition to inversion, would generate high enough ankle ligament
strain in the ATaFL to cause a grade I injury.

METHODS:
The development and validation of a three-dimensional multibody dynamic foot model to
study the effect of foot constraint on ankle injury due to external rotation have been
described in detail by Wei et al. [7], thus only a brief description is given here. Joint
anatomical features were taken from a CT set. Detailed features of the foot/ankle
complex were obtained by importing Digital Imaging and Communications in Medicine
(DICOM) files from an individual CT scan into Materialise’s Interactive Medical
Imaging Control System (MIMICS) (Materialise, Ann Arbor, MI). This yielded a threedimensional surface model of the bones as Stereolithography (STL) files for export. To
reduce the size of the surface files and subsequent model, the STL files were remeshed in
MIMICS to smooth the surface of each bone. Exported files were then imported into the
three-dimensional solid modeling software SolidWorks (TriMech Solutions, LLC,

71

Columbia, MD) as Mesh Files. SolidWorks along with the ScanTo3D package was used
to further construct each bone and simplify the bone surfaces. SolidWorks Motion was
then used to assemble bones, obtain proper positioning, add necessary components, and
run simulations. The tibia was fixed in space, the fibula, talus, and calcaneus were free to
move, and the rest of the bones were fused together and moved as a rigid body. Twenty
ligaments were included and represented as linear spring elements (Fig 4-1), with
stiffness values from the literature [7]. A model sensitivity analysis [7] was conducted to
ensure that moderate ligament stiffness variations do not significantly alter conclusions
drawn from the model.

To simulate the injury event documented by Fong et al. [3], approximately 3 times body
load (1840 N) was applied to the proximal end of the model, proportionally distributing it
between the tibia and the fibula with one-sixth of the loading in the fibula [7]. Fong et al.
[3] documented 48° of ankle inversion at injury. In addition, from a motion analysis by a
skeleton model-based image-matching technique they also reported a maximum 20° of
plantar flexion and 30° of internal rotation. To separate the motions and investigate
mechanisms of the injury, three motion scenarios were simulated in this study: (1) a pure
inversion of 48°, with an axis of rotation passing through the calcaneus anteroposteriorly;
(2) a 48° inversion and at the same time a 20° plantar flexion, with an axis of rotation
passing through the talus medial-laterally; (3) the motions in (2) and simultaneously a 30°
internal rotation, with an axis of rotation along the tibia and passing through the talus.
Ligament strains were determined in the model.

72

Figure 4-1 Schematic of the generic ankle model showing the locations of 20 modeled
ligaments: the interosseous ligaments (Inte-I and Inte-II); the anterior and posterior
tibiofibular (ATiFL and PTiFL); the calcaneofibular (CaFL); the anterior and posterior
talofibular (ATaFL and PTaFL); the anterior and posterior tibiotalar (ATiTL and PTiTL);
the tibionavicular (TiNL); the talonavicular (TaNL); the interosseous, lateral, medial, and
posterior talocalcaneal (ITaCL, LTaCL, MTaCL, and PTaCL); the calcaneocuboid
(CaCL); the calcaneonavicular (CaNL); the medial, central, and lateral plantar fascia
(MPF, CPF, and LPF).

73

RESULTS:
Comparisons of different motion scenarios to the ankle neutral position were
demonstrated in Fig 4-2. With the ankle in inversion, plantar flexion, and internal rotation,
Fig 4-2d showed the largest ankle motion and possibly the maximum ankle instability.
(a)

(b)

(c)

(d)

Figure 4-2 Posterior views of the left ankle model in (a) neutral position; (b) pure
inversion; (c) inversion and plantar flexion; (d) inversion, plantar flexion, and internal
rotation. For clarity, ligaments were not shown.

Ligament strains were compared for the three motion scenarios (Fig 4-3). For pure
inversion, the CaFL strained the most at 12%, and the ATaFL and the LTaCL were in
lower at about 10%. For the scenario of Inv.+ Plan.Flex.+ Int.Rot., the highest strain
occurred in the ATaFL (20.5%), followed by the CaFL (16%).

74

25
Inv.

Strain (%)

20

Inv.+ Plan.Flex.
Inv.+ Plan.Flex.+ Int.Rot.

15
10
5

aC
L
PT

L
IT
aC

C
L
LT
a

aF
L
PT

Ta
FL
A

aF
L
C

iF
L
PT

A

Ti
FL

0

Ligaments

Figure 4-3 Strains in various ligaments for different ankle motions. Eight ligaments were
selected based on strain being higher than 2% for at least one motion.

DISCUSSION:
This study supported the hypothesis that plantar flexion and internal rotation would play
important roles in an ankle inversion injury. The highest strain seen in the ATaFL with a
combination of inversion, plantar flexion, and internal rotation is in concert with the
motions documented in the accidental ankle injury noted by Fong et al. [3]. While
clinically the most common mechanism of injury in a lateral ankle sprain is a
combination of inversion and plantar flexion [10], those conclusions are drawn mostly
from patients’ questionnaires and may not reflect the actual ankle joint biomechanics
during injury [3]. This simulation study and the injury case study by Fong et al. [3] may
reveal the importance of internal rotation, add to the knowledge about the potential
mechanisms of ankle sprain, and raise debate on ankle joint orientation during an
inversion sprain injury.
75

Fong et al. [3] indicated that the injury event was a grade I mild ATaFL sprain. Within
the literature, there is a lack of correlation between ankle sprain severity and ankle
ligament strains. Sprains of ankle ligaments range in severity from grade I to grade III [6].
The grade I ligament sprain is characterized by minimum tearing of ligament fibers [6].
Knowing that failure strain of ankle ligaments is in the range of 30-35% [4], this
simulation study suggests a grade I sprain may correlate with ligament strains of 15-20%
(Fig 4-3), which is supported by the fact that ligament collagen fibers start tearing after
50% of failure strain [11]. A grade III ankle ligament sprain is described as rupture of a
ligament [6], therefore, based on the current simulation, a grade II sprain should parallel
with ligament strains of 20-30%.

In conclusion, the ability of the computational model to predict a low level of ankle
ligament sprains is encouraging. The model may be used in future simulations to provide
a computational basis for studies of other ankle injuries or ankle sprain prevention
strategies.

76

PART II: MECHANISM OF INJURY IN A HIGH ANKLE SPRAIN

INTRODUCTION:
Injury to the tibiofibular syndesmosis ligaments, which bind together the distal ends of
the tibia and fibula, is commonly referred to as a high ankle sprain [10]. While lateral
ankle sprains are the most common injury, high ankle sprains represent a more disabling
problem and require a longer recovery period [1] and different treatment [5]. The
mechanism associated with a high ankle sprain is primarily thought to involve external
rotation of the foot [1,9]. However, both a cadaver study [8] and a simulation study [7]
show that tibiofibular syndesmosis ligaments are not stretched the most during an
excessive, pure external foot rotation.

Sports-related injury mechanisms can be studied through various approaches: cadaveric
experiments, kinematics, biomechanics, and computational modeling. While cadaver
studies may not reflect ligament strains during injury-causing events, in vivo
experimental studies of ligament rupture experienced during sports activities are not
feasible with human subjects. Dynamic three-dimensional simulation studies offer an
attractive alternative because of their ability to separate motions and study them
individually, rapidly solve for motion-based mechanics, and determine ligament strains
during large motions.

Recently, three injury cases were reported and diagnosed as high ankle sprains in NCAA
football games. A review of the game films indicated that all three injuries involved
dorsiflexion and eversion of the foot by a blow to the lateral aspect of the shank and knee,
77

while the player’s foot was planted in external rotation. The purpose of this study was to
test the influences of each of these various motions on tibiofibular syndesmosis ligament
strains using a multibody three-dimensional foot model, driven by the game film
examination of the injury events. The hypothesis was that a combination of external
rotation, dorsiflexion, and eversion applied to the ankle would generate high strain in the
anterior tibiofibular ligament (ATiFL) to form a basis for the diagnosed high ankle sprain.

METHODS:
A three-dimensional multibody foot model was developed based on computed
tomography (CT) scans and validated for measurements of ankle ligament strain [2] and
ankle joint torque [8] against two well-documented cadaver studies. The model has
shown its ability to predict ankle injuries due to excessive levels of external foot rotation
[7]. In addition to a pure external rotation, which simulates the previous cadaver
experiments, dorsiflexion and/or eversion were also applied to this foot model to
investigate the influences of these motions on the behavior of ATiFL.

The development and validation of this foot model to predict ankle ligament strains have
been described in detail by Wei et al. [7], so only a brief description will be given here.
Joint anatomical features were taken from a generic CT set. Detailed features of the
foot/ankle complex were obtained by importing Digital Imaging and Communications in
Medicine (DICOM) files from an individual CT scan into Materialise’s Interactive
Medical Imaging Control System (MIMICS) (Materialise, Ann Arbor, MI). This yielded
a three-dimensional surface model of the bones as Stereolithography (STL) files for

78

export. To reduce the size of the surface files and subsequent model, the STL files were
remeshed in MIMICS to smooth the surface of each bone. Exported files were then
imported into the three-dimensional solid modeling software SolidWorks (TriMech
Solutions, LLC, Columbia, MD) as Mesh Files. SolidWorks along with the ScanTo3D
package was used to further construct each bone and simplify the bone surfaces.
SolidWorks Motion was then used to assemble bones, obtain proper positioning, add
necessary components, and run the simulations. The tibia was fixed in space, the fibula,
talus, and calcaneus were free to move, and the rest of the bones were fused together and
moved as a rigid body. Twenty ligaments were included and represented as linear spring
elements (Fig 4-1), with stiffness values from the literature [7]. A model sensitivity
analysis [7] was conducted to ensure that moderate ligament stiffness variations did not
significantly alter conclusions drawn from the model.

Approximately 3 times body weight (2000 N) was applied to the proximal end of the
model, proportionally distributing it between the tibia and the fibula with one-sixth of the
loading in the fibula [7]. To separate the motions and investigate mechanisms of injury,
four motion scenarios were simulated in this study: (1) a pure external rotation of 30°,
with an axis of rotation along the tibia and passing through the talus; (2) a 30° external
rotation and a 20° dorsiflexion, with an axis of rotation passing through the talus mediallaterally; (3) a 30° external rotation and a 20° eversion, with an axis of rotation passing
through the calcaneus anteroposteriorly; (4) a combined 30° external rotation, 20°
dorsiflexion, and 20° eversion. Ligament strains were determined in the model.

79

RESULTS:
Comparisons of different motion scenarios with the neutral ankle are demonstrated in Fig
4-4.

(a)

(b)

(c)

(d)

(e)

Figure 4-4 Anterior views of the left ankle model in (a) neutral position; (b) pure
external rotation; (c) external rotation and dorsiflexion; (d) external rotation and eversion;
(e) external rotation, dorsiflexion, and eversion.

Ligament strains were compared for combinations of the four motion scenarios (Fig 4-5).
Under a pure external rotation, the ATiTL strained the most at 8%, while the PTaFL and
the ATiFL were strained about 6% and 3.5%, respectively. Dorsiflexion and eversion,
added to the external rotation, significantly increased ATiFL strain to 9% and 17%,
respectively. For the combination of all three motions, the highest strain occurred in the
ATiFL (21%), followed by the ATiTL (15.5%).

80

25

Ext.Rot.
Ext.Rot.+ Dor.Flex.
Ext.Rot.+ Eve.
Ext.Rot.+ Dor.Flex.+ Eve.

Strain (%)

20
15
10
5
0
Inte-II

ATiFL

PTiFL

ATiTL

TiNL

PTaFL

Ligaments

Figure 4-5 Strains in various ligaments for different ankle motions. Six ligaments were
selected based on strain being higher than 5% for at least one motion.

DISCUSSION:
This study quantitatively showed that external rotation, dorsiflexion and eversion each
played an important role in a high ankle sprain. The highest strain observed in the ATiFL
occurred with a combination of all three motions, in concert with the motions seen in
game films. The results are supported by clinical findings that syndesmosis sprains
involve excessive dorsiflexion and eversion of the ankle joint with external rotation of the
foot [10].

A limitation of the study was that from the game films themselves it was impossible to
accurately measure foot and ankle rotations during the injury, which may affect input
simulations and consequently the accuracy of the ligament strains. A recent study, using
the same foot model, simulates a well-documented, laboratory accidental ankle sprain
81

event [3]. It was concluded that a grade I ankle sprain may parallel with ligament strains
of 15-20%. Thus, with input motions of 30° external rotation, 20° dorsiflexion, and 20°
eversion, the current study suggested a grade I level of ligament sprain (Fig 4-5). Future
studies, that better incorporate injury video data, may provide a more accurate correlation
between these types of clinical injuries and ligament strains.

In conclusion, the ability of the computational model to predict high strain in the ATiFL
is encouraging. The model may be used in future simulations to provide a computational
basis for studies of other ankle injuries and help develop prevention strategies.

82

REFERENCES

83

REFERENCES

1. Boytim MJ, Fischer DA, Neumann L. Syndesmotic ankle sprains. Am J Sports
Med. 1991;19(3):294-298.
2. Colville MR, Marder RA, Boyle JJ, Zarins B. Strain measurement in lateral ankle
ligaments. Am J Sports Med. 1990;18(2):196-200.
3. Fong DT, Hong Y, Shima Y, Krosshaug T, Yung PS, Chan KM. Biomechanics of
supination ankle sprain: a case report of an accidental injury event in the
laboratory. Am J Sports Med. 2009;37(4):822-827.
4. Funk JR, Hall GW, Crandall JR, Pilkey WD. Linear and quasi-linear viscoelastic
characterization of ankle ligaments. J Biomech Eng. 2000;122(1):15-22.
5. Hopkinson WJ, St Pierre P, Ryan JB, Wheeler JH. Syndesmosis sprains of the
ankle. Foot Ankle. 1990;10(6):325-330.
6. Noyes FR, Grood ES, Torzilli PA. Current concepts review. The definitions of
terms for motion and position of the knee and injuries of the ligaments. J Bone
Joint Surg Am. 1989;71(3):465-472.
7. Wei F, Hunley SC, Powell JW, Haut RC. Development and validation of a
computational model to study the effect of foot constraint on ankle injury due to
external rotation. Ann Biomed Eng. 2011;39(2):756-765.
8. Wei F, Villwock MR, Meyer EG, Powell JW, Haut RC. A biomechanical
investigation of ankle injury under excessive external foot rotation in the human
cadaver. J Biomech Eng. 2010;132(9):091001.
9. Williams GN, Jones MH, Amendola A. Syndesmotic ankle sprains in athletes. Am
J Sports Med. 2007;35(7):1197-1207.
10. Wolfe MW, Uhl TL, Mattacola CG, McCluskey LC. Management of ankle
sprains. Am Fam Physician. 2001;63(1):93-104.
11. Yahia L, Brunet J, Labelle S, Rivard CH. A scanning electron microscopic study
of rabbit ligaments under strain. Matrix. 1990;10(1):58-64.

84

CHAPTER 5
EVERSION DURING EXTERNAL ROTATION OF THE HUMAN CADAVER FOOT
PRODUCES HIGH ANKLE SPRAINS

ABSTRACT
While high ankle sprains are often clinically ascribed to excessive external foot rotation,
no experimental study documents isolated anterior tibiofibular ligament (ATiFL) injury
under this loading. This in vitro study hypothesized that external rotation of a highly
everted foot would generate ATiFL injury, in contrast to deltoid ligament injury from
external rotation of a neutral foot. Twelve (six pairs) male cadaveric lower extremity
limbs underwent external foot rotation until gross failure. All limbs were positioned in 20
degrees of dorsiflexion and restrained with elastic athletic tape. Right limbs were in
neutral while left limbs were everted 20 degrees. Talus motion relative to the tibia was
measured using motion capture. Rotation at failure for everted limbs (46.8 ± 6.1 deg) was
significantly greater than for neutral limbs (37.7 ± 5.4 deg). Everted limbs showed ATiFL
injury only, while neutral limbs mostly demonstrated deltoid ligament failure. This is the
first biomechanical study to produce isolated ATiFL injury under external foot rotation.
Eversion of the axially loaded foot predisposes the ATiFL to injury, forming a basis for
high ankle sprain. The study helps clarify a mechanism of high ankle sprain and may
heighten clinical awareness of isolated ATiFL injury in cases of foot eversion prior to
external rotation. It may also provide some guidance to investigate the effect of
prophylactic measures for this injury.
Keywords: biomechanical study; syndesmotic ankle sprain; external rotation; talus
motion; anterior tibiofibular ligament.

85

INTRODUCTION
Injury to the distal tibiofibular syndesmotic ligaments is commonly referred to as a high
ankle sprain or syndesmotic ankle sprain. The distal tibiofibular syndesmosis is
composed of the anterior and posterior tibiofibular ligaments (ATiFL and PTiFL) and the
interosseous ligament (IOL).1 The ATiFL is the main structure of the ankle joint
preventing excessive fibular movement and external rotation of the talus.2 Among the
ligaments comprising the distal tibiofibular syndesmosis, the ATiFL is the most
susceptible soft tissue to high ankle sprains.25 High ankle sprains are less common than
lateral ankle sprains, comprising approximately 10% of all ankle sprains.3,4 However,
they represent a more disabling problem5 requiring a longer recovery period,3 different
treatment4 and can lead to chronic ankle dysfunction.6 While high ankle sprains have
historically been underdiagnosed, more recently a heightened awareness in the sports
medicine community has resulted in more frequent diagnoses of high ankle sprains.7

In the clinical literature, the mechanism of high ankle sprain has mostly been attributed to
external foot rotation.3,4,8-10 A study that reviews high ankle sprains in athletes, however,
has indicated that the mechanism of injury involves “the ankle being subjected to an
external rotation moment with the foot in a dorsiflexed, pronated position”.7 This
combined dorsiflexion, eversion, and external rotation mechanism has been supported
clinically by others.5,11-13

Numerous biomechanical studies have been conducted to help understand mechanisms of
high ankle sprain in controlled laboratory experiments. Many of these, however, have

86

been ligament sectioning studies that show increased lateral motion of the fibula relative
to the tibia during external rotation of the foot as syndesmotic structures are sequentially
cut.14-18 Other biomechanical studies have largely been designed to investigate ankle
fractures rather than sprains.19,20 For example, a study by Haraguchi and Armiger,21 using
a relatively rigid experimental restraint for the foot, has shown external rotation of the
foot with the ankle in pronation produces damage to the ATiFL with fibular fracture,
followed by medial injury that involves the deltoid ligament. Another early study by
Stiehl et al.,22 using a more compliant fiberglass cast tape restraint of the foot onto a rigid
plate, documents ATiFL injury in combination with fibular fracture. While classically a
patient with a high ankle sprain reports tenderness over the anterolateral tibiofibular joint
where the fibers of the ATiFL are often disrupted,33 a review of the literature by the
current authors found no previous in vitro studies producing isolated ligamentous injuries
to the ankle that represent a clinical high ankle sprain.

Wei et al.23 recently developed a computational model of the ankle to simulate the effects
of experimental restraint on the development of strains in various ankle ligaments during
external rotation of the foot based on laboratory experiments by Villwock et al.24 The
model analysis showed the largest strains developing in the anterior deltoid ligament
(ADL), rather than the ATiFL, for 30 degrees of foot rotation. The model has also been
recently utilized in a study investigating ankle ligament strains during simulation of ankle
motions documented based on game films during high ankle sprain injuries on the
football field.25 The films suggested that these injuries involved dorsiflexion and eversion
of the foot from blows to the lateral aspect of the shank and knee, while the player’s foot

87

was planted in external rotation. The model analysis showed that the eversion component
of talus motion for an axially loaded foot played a very significant role in developing
large strains in the ATiFL during external foot rotation, while strains in the ADL, PTiFL,
and IOL were relatively low. These large strains may predispose the ATiFL to injury.

As a result of these earlier studies by our laboratory, the purpose of the current study was
to conduct experiments using human cadaver ankles, as guided by the computational
model, for investigations of external rotation of a highly everted foot. It was hypothesized
that for an axially loaded, dorsiflexed ankle, eversion of the foot prior to application of
external rotation would generate injury to the ATiFL, in contrast to the ADL injury for a
neutral foot condition.

METHODS
Specimen Preparation
The experiments were conducted on twelve (six pairs) fresh-frozen lower limbs from
male cadavers (aged 56 ± 12 years). The limbs were stored at -20°C and thawed to room
temperature for 24 hours before tests. The tibia and fibula were transected approximately
15 cm distal to the center of the knee. The proximal end of the tibia and fibula shafts
were then cleaned with 70% alcohol and potted in an aluminum box with room
temperature curing epoxy (Fiber Strand, Martin Senior Corp., Cleveland, OH) (Figure 51a). Two screws were placed in the medial and lateral aspects of the proximal tibia, with
an approximately 30 mm projected length, to help restrain the tibia from rotating within
the potting material (Figure 5-1b). Athletic elastic tape (Elastikon, Johnson & Johnson,

88

New Brunswick, NJ) was used to restrain the foot onto a polycarbonate plate (22 x 10 cm)
(Figure 5-1c). Moderate tension was placed on the tape during this process. The plate was
then inserted into a rectangular aluminum tray (Figure 5-1d). Both the left and right limbs
were dorsiflexed 20 degrees by means of a rigid wedge placed under the foot plate
(Figure 5-1d). The aluminum tray with the left limbs was then manually rotated and
locked with 20 degrees of eversion in the coronal plane. The right limbs were in a neutral
position. The specimen was mounted upside down (foot pointing upward) in the test
machine (Figure 5-1e).

89

b

a

d

c

screw
locker

foot plate

aluminum
tray

20 deg wedge

e

talus & tibia
marker sets

aluminum
box

Figure 5-1 Specimen preparation and setup. The proximal end of the shank was potted
using epoxy resin (a) with two screws placed earlier into the proximal tibia (b). The foot
was restrained onto a polycarbonate plate by athletic elastic tape (c). A custom fixture (d)
had an aluminum tray which was able to rotate in the coronal plane, providing eversion of
the foot. The limb with markers was mounted upside down into the testing machine (e).
90

Testing Protocol
Experiments were conducted at room temperature on a custom, hydraulic, biaxial testing
machine using a 244 Nm rotary actuator (Model SS-001-1V, Micromatic, Berne IN) and
a vertically oriented linear actuator (Model 204.52, MTS Corp.) The aluminum box
(Figure 5-1e) was attached to the rotary actuator through a biaxial (torsion-axial) load cell
(Model 1216CEW-2K, Interface, Scottsdale, AZ) with a capacity of 2000 lbs (8896 N)
axial force and 1000 lbs-in (113 Nm) torsion (Figure 5-2). The foot was secured to the
rotation-locked linear actuator with a custom fixture that allowed x-y adjustments to align
the tibia along the linear and torsional actuators. A compressive preload of 1500 N,
approximately two times body weight, was applied axially to simulate weight bearing in a
dynamic situation.28 Internal tibia rotation (external foot rotation) was input in position
control at a frequency of 1 Hz (0.5 s to peak). This rotation rate was similar to that
documented for real injury events.31 The tests were started with 20 degrees of rotation
and repeated in increments of 10 degrees successively until a diagnosed injury in the
ankle, via the torque signal. The starting point of rotation and 10 degree increments were
selected to ensure failure within a minimal number of tests. Approximately five minutes
were allowed between successive increments for the investigators to assess damage to the
ankle, review the collected data, and assure that the orientation of the foot to the plate
was not altered between tests. The mode of injury, failure torque, and foot rotation at
failure were documented for each specimen. Following the tests, a careful gross
dissection of the joint was performed by an orthopaedic surgical resident (JMP) to
document all soft tissue injuries.

91

1X video
camera

5X Vicon
cameras

biaxial
load
cell

Figure 5-2 Testing setup. Five-camera Vicon motion capture system and one video
camera were used to track motions of the talus relative to the tibia.

Motion Analysis
All specimens were subjected to a pre-test computed tomography (CT) scan performed
with a reflective marker array screwed into the talus. This array was utilized in
subsequent analyses of talus motion with respect to the tibia using a Vicon motion
capture system (Oxford Metrics Ltd., Oxford, United Kingdom) (Figure 5-2). The talus
array was attached from the posteromedial aspect of the ankle and care was taken to
avoid damage to the posterior deltoid ligament or the PDL (posterior tibiotalar and
tibiocalcaneal ligaments). The tibia array was positioned 10-20 cm proximal to its
inferior articular surface (Figure 5-1e). A joint coordinate system (JCS) was established
92

based on motion of the reflective marker array, as described in a previous study.23 The
medial-lateral translation and rotation of the talus relative to the tibia were determined in
this JCS for the 30 degrees of foot rotation test on each limb.

Statistical Analysis
One-way ANOVA and Student-Newman-Keuls (SNK) post hoc tests were used to
statistically compare the differences in failure torque, failure rotation, and talus motion
(translation and rotation) relative to the tibia between the paired limbs. Two-way
ANOVA and SNK post hoc tests were used to determine the differences in the torquerotation data between limbs (factor one) at each rotation level (factor two). In all
statistical tests p values less than 0.05 were considered significant.

RESULTS
The torque-rotation responses were significantly different between limbs, with the neutral
foot stiffer than the everted one (Figure 5-3). Tissue failure was suspected when a distinct
drop in the torque-time curve was observed (Figure 5-4a). The torque-rotation curves
exhibited loading and unloading parts (hysteresis loops) that indicated energy loss as a
result of ligamentous rupture (Figure 5-4b). While failure torque for the highly everted
foot was not significantly different than for the neutral foot (90.3 ± 8.6 Nm versus 84.4 ±
16.9 Nm), failure rotation for the highly everted foot (46.8 ± 6.1 deg) was significantly
greater than for the neutral foot (37.7 ± 5.4 deg) (Table 5-1). The injury modes for the
neutral limbs were: one partial tear of the ADL, one rupture of the ADL, one rupture of
the deltoid ligament, one tibial avulsion of the ADL, one combined partial tear of the

93

ADL and ATiFL, and one fibular fracture. The injury modes for the everted limbs were:
four total ruptures of the ATiFL and two partial tears of the ATiFL (Table 5-1). Typical
gross injuries were shown in Figure 5-5.

100
80

Torque (Nm)

§

neutral
everted

‡

60

‡
‡

40
‡

20
0
0

10

20

30

40

External Rotation (deg)
Figure 5-3 Torque-rotation curves of different limbs were constructed from the
subfailure test data. § indicates significant difference between curves. ‡ indicates
significant difference between limbs at various rotation points.

During axial loading of the ankle and prior to external rotation, lateral translation of the
talus was significantly greater in everted limbs (2.67 ± 0.74 mm) than neutral limbs (0.72
± 0.57 mm) (Table 5-2). Talus eversion during axial loading of the ankle was not
different between limbs. During 30 degrees of external foot rotation, the talus externally
rotated more in everted limbs (26.8 ± 1.4 deg) than neutral limbs (18.3 ± 1.8 deg), and

94

lateral translation of the talus was also significantly greater for everted limbs (2.88 ± 1.13
mm) than neutral limbs (0.12 ± 0.95 mm).

100

a

subfailure
failure

Torque (Nm)

80
60
40
20
0
0

0.5

1

1.5

2

2.5

3

-20
Time (s)
100

b

subfailure
failure

Torque (Nm)

80
60
40
20
0
0

10

20

30

40

50

-20
External Rotation (deg)
Figure 5-4 Typical temporal profile of torque for subfailure and failure tests (a). The
torque-rotation curves had loading and unloading parts (b). The subfailure curve was one
increment (10 deg) prior to the failure curve. Data were from the neutral limb of
specimen 1.
95

Table 5-1 Specimen descriptions, torque and foot rotation at failure, and resultant injury.

Specimen

Height Weight
Age
(m)
(kg)

Failure torque
(Nm)

Failure rotation
(deg)

Neutral

Everted

Neutral

Everted

1

76

1.72

64

87.4

78.3

36.0

43.4

2

56

1.83

79

80.9

91.8

47.1

55.3

3

40

1.88

75

88.0

94.5

35.1

50.1

4

50

1.75

88

101.5

103.8

39.8

50.0

5

55

1.78

66

53.1

87.2

31.1

43.4

6

56

1.88

93

95.3

86.5

37.0

38.6

Mean

56

1.81

77.5

84.4

90.3

37.7

46.8*

SD

12

0.07

11.6

16.9

8.6

5.4

6.1

Failure mode
Neutral
Tibial avulsion of
the ADL
Partial tear of the
ADL
Partial tear of the
ADL and ATiFL
Rupture of the
deltoid ligament
Spiral fracture of
the fibula
Rupture of the
ADL

Everted
Rupture of the ATiFL
Partial tear of the ATiFL
Rupture of the ATiFL
Rupture of the ATiFL
Partial tear of the ATiFL
Rupture of the ATiFL

The posterior tibiofibular ligament and the
interosseous ligament remained intact in all
specimens.

Note: The right limbs were neutral in the coronal plane, while the left limbs were everted 20°.
ATiFL is the anterior tibiofibular ligament.
ADL is the anterior deltoid ligament, which includes the anterior tibiotalar and tibionavicular ligaments.
Deltoid ligament involves the ADL, the posterior tibiotalar ligament, and the tibiocalcaneal ligament.
* Statistically different than the neutral foot.

96

b

a

Fibula
Fibula
Tibia

Tibia

c

d

Tibia

Tibia

Figure 5-5 Typical gross injuries. Rupture of the ATiFL (a); Partial tear of the ATiFL
(b); Rupture of the deltoid ligament (c); Rupture of the ADL (d). Arrows indicate the
sites of injury.

Table 5-2 Talus motion relative to the tibia (mean ± 1SD).
Axial loading of 1500 N

External foot rotation of 30°

Talus motion
Neutral
Lateral translation (mm)
External rotation (deg)
Eversion (deg)

Everted

Neutral

Everted

0.72 ± 0.57

2.67 ± 0.74*

0.12 ± 0.95

2.88 ±1.13*

---

---

18.3 ± 1.8

26.8 ± 1.4*

5.0 ± 1.8

3.5 ± 2.3

---

---

Note: * Statistically different than the neutral foot (p<0.001).
--- Negligible.

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DISCUSSION
Results of the current study supported the hypothesis that external rotation of the highly
everted foot would generate a high incidence (6 of 6 limbs) of isolated injury to the
ATiFL. In contrast, external rotation of the neutral foot (in the coronal plane) produced
damage to the deltoid ligament in 5 of 6 limbs. This is the first biomechanical study to
produce isolated ATiFL injury using intact cadaver limbs in a laboratory. Foot eversion
was found to separate high ankle sprain from deltoid ligament injury. This implied that
by everting the foot, external rotation force may be transferred from the ADL to the
ATiFL. Lateral translation and external rotation of the talus were found to be
significantly greater in the everted limbs than in the neutral ones.

The mechanism of injury for high ankle sprains can be confusing because of the different
anatomic structures involved and the manner in which these structures develop excessive
stress in the three planes of motion.26 Knowing the role of the ATiFL in talus rotation is
important in understanding the mechanism of injury in a high ankle sprain. In addition to
holding the fibula tight to the tibia, this ligament prevents excessive movement of the
fibula and external rotation of the talus.2 In the current study, motion analysis of the talus
relative to the tibia helped explain the mechanism of injury. In a highly everted foot, axial
loading caused lateral translation of the talus (Figure 5-6b), thereby pushing the fibula
away from the tibia, and potentially producing pre-strain of the ATiFL, as this ligament is
oriented medial-laterally in the coronal plane. During external foot rotation the talus was
forced to rotate externally and translate laterally, resulting in further separation of the
fibula from the tibia causing disruption of the ATiFL (Figure 5-6c). In the neutral foot,

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however, lateral translation of the talus was minimal during both axial loading and
external rotation (Table 5-2). Consequently, external rotation of the talus resulted in ADL
rupture, as this ligament has an antero-posterior direction in the sagittal plane (Figure 56f). This mechanism of ADL injury has been supported in previous studies. A clinical
review indicates that external rotation of the neutral foot will “first rupture the deltoid
ligament, with subsequent injury to the ATiFL”.1 Using a generic computational ankle
model, a simulation study by our laboratory also shows that the ADL experiences the
highest strain under external rotation of the neutral foot with a compliant foot
constraint.23 In addition to the mechanism of ADL injury associated with axial rotation,
other mechanisms potentially causing deltoid ligament damage could be excessive
eversion, dorsiflexion, or a combination of both.19 However, these were not the subject of
the current study.

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a

b

c

fibula

tibia

ATiFL

talus

d

tibial
plafond

e

f

talar
dome

ADL

Figure 5-6 Schematics of a left ankle joint in an anterior view (a b c d) and a medial view
(e f). The ATiFL is shown in (a b c) with an everted talus, and the ADL is shown in (d e
f) with a neutral talus. The ATiFL has a medial-lateral direction with an oblique angle of
~30° lateral-distally. The ADL is oriented anteroposteriorly in the sagittal plane. (a),
unloaded ankle. (b), when axial loading was applied, the talus translated laterally. (c),
when external rotation was applied, the talus further translated laterally, resulting in
ATiFL rupture. (d and e), unloaded ankle in anterior and medial views. (f), when axial
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Figure 5-6 (cont’d)
loading and external rotation were applied to the neutral ankle, the talus externally
rotated, resulting in ADL rupture.

Results of the current study showed that, compared to neutral limbs, everted limbs were
less stiff (Figure 5-3) and, as a result, experienced more rotation prior to gross failure of
the ankle (Table 5-1). One possible reason may be that the talus in a neutral foot is fully
restrained by the ankle mortise, which limits its external rotation during rotation of the
foot (Figure 5-6d). When an axial load is applied to a dorsiflexed foot (neutral in the
coronal plane), the tibial plafond tends to hold the talus tightly in the ankle mortise,
resulting in large ankle joint forces being generated during external rotation of the talus.
In fact, an excessive axial loading of a dorsiflexed ankle often produces fractures of the
distal tibia or the talus.19 In contrast, an angle is formed between the tibial plafond and
the talar dome in an everted foot, causing the talus to move laterally during axial loading,
perturbing the ankle mortise structure, resulting in its ease of rotation. This was supported
by the results shown in Table 5-2 in which external rotation of the talus was significantly
greater for the everted limbs than for the neutral limbs.

The neutral and everted limbs failed at comparable levels of torque, but significantly
different levels of foot rotation (Table 5-1). The ankle with a neutral foot failed at smaller
input rotations. Research by Lundberg et al.27 suggests that rotations of the bones within
the foot during external rotation may act as torsion dissipating devices, implying that a
rotation dependent injury mechanism may be influenced by compliance of the ankle joint
structure. The everted limbs in the current study were more compliant than the neutral

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limbs. But the fact that both limbs failed at the same torque level suggests that ankle joint
torque may be a crucial injury risk factor directly related to ankle ligament strains. And
yet, the location of ligamentous injury at a given torque was shown to be highly
dependent on foot orientation in this study.

Some limitations of the study should be noted. The effect of transection of the tibia and
fibula, resulting in a lack of the proximal tibiofibular syndesmosis in the testing
specimens, on potential alterations of ankle joint mechanics was unknown. And, the
repeated testing methodology may have generated micro-damages to soft tissues and
other joint structures before the failure test. However, similarity in the loading portion of
successive tests, as evidenced by the typical torque-rotation curves between failure and
subfailure tests (Figure 5-4b), may indicate negligible micro-damages to the soft tissues
during this procedure. An advantage of the repeated testing procedure, however, was the
ability to control the extent of damage to the ankle. Previous studies, by others, have
described combinations of bone fracture and ligamentous injury with difficulty in
clarifying which injury occurs first.21,22 Another limitation of the current study was that a
constant frequency of 1 Hz was used in the foot rotation tests. Since the angle of rotation
was different between tests, the rate of rotation varied. However, a previous cadaver
study28 shows variations in rotation rate do not significantly affect the stiffness of
foot/ankle complex or the mechanism of ankle failure. A common limitation of cadaver
studies is that the effects of active muscle are not included. A recent study reveals a 1020% difference in ankle joint torques between tests of human subjects and cadavers, with
living subjects generating higher torques.29 While the effects of active muscle on talus

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motion and subsequent ankle injury occurring in sports are currently unknown, kinematic
data may be generated in future studies using game films of injury-producing events.30,31
Computational models may then be used to estimate soft tissue (muscle, ligaments)
deformations.23,29 Specimen age may also by a factor that affects ligament properties.
While the study was not conducted on a typical young, sport-active population,
contralateral limbs of each subject were used to study the effect of foot eversion. The
assumption here was that no differences in ankles or their properties existed between
limbs, as the authors were not aware of any current literature suggesting otherwise.
Furthermore, while slight differences in specimen length may have existed between
subjects in the study, axial load would not be significantly affected. And, since injuries
were consistent between different states of eversion, the issue of specimen length was not
indicated to be a significant limitation of the current study. And finally, while tension in
the commercial athletic tape could have varied slightly between tests, the tape restraint of
the foot used in this study may have allowed more motion of the bones in the foot than
more rigid experimental foot restraints used in previous biomechanical studies to produce
soft tissue injuries rather than bone fractures. The protocol to control foot rotation
magnitudes incrementally between tests may have also helped limit injury to soft tissue
structures first. Future studies could be conducted to continue the incremental loading
protocol and determine if subsequently bone fracture might ensue.

Other future studies can be envisioned. Firstly, the current study showed that 20 degrees
of foot eversion transferred injury from the ADL to the ATiFL. Simulation studies, using
computational models, could be conducted in the future to parametrically vary levels of

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talus eversion prior to external rotation, and quantify ankle ligament strains to determine
the potential threshold of eversion that causes a switch in the location of injury. Secondly,
both the cemented foot restraint of earlier studies21,28 and the more flexible foot restraint
to a base plate of the current study may not be comparable to the foot restraint in a sports
setting. Future cadaver studies need to involve shoes and investigate the influence of shoe
restraint in the location of ankle injury. Finally, lateral translation of the talus due to
eversion of the axially loaded foot, as evidenced by the current in vitro investigation, may
be studied in vivo in a weight bearing ankle using fluoroscopic techniques32 or an upright
MRI.

In conclusion, this biomechanical study restrained pairs of cadaver ankles using athletic
tape, externally rotated axially pre-loaded limbs in dorsiflexion with and without foot
eversion, and identified isolated ligamentous injury to the ATiFL in everted limbs. The
resultant high incidence of ligamentous injuries (11 of 12 ankles), as opposed to bone
fracture in other biomechanical studies, implied that the taping technique may be
appropriate for future laboratory researchers to study ankle soft tissue injuries common in
sports. This is the first biomechanical study to produce isolated ATiFL injury under
external rotation of a highly everted foot. Eversion of an axially loaded foot predisposes
the ATiFL to injury which may describe one clinical basis for high ankle sprains. Despite
its biomechanical nature, the current study may heighten clinical awareness of isolated
ATiFL injury in cases of foot eversion and dorsiflexion prior to external rotation, as
opposed to deltoid ligament injury in cases of foot dorsiflexion followed by external
rotation, and provide some guidance to help study the effect of prophylactic measures for

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high ankle sprains, such as wedged shoe insoles or ankle braces to help limit excessive
foot eversion. However, interventions may need to be balanced with altered player
performance issues.

105

REFERENCES

106

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CHAPTER 6
ROTATIONAL STIFFNESS OF FOOTBALL SHOES INFLUENCES TALUS
MOTION DURING EXTERNAL ROTATION OF THE FOOT

ABSTRACT
Shoe-surface interface characteristics have been implicated in the high incidence of ankle
injuries suffered by athletes. Yet, the differences in rotational stiffness among shoes may
also influence injury risk. It was hypothesized that shoes with different rotational
stiffness will generate different patterns of ankle ligament strain. Four football shoe
designs were tested and compared in terms of rotational stiffness. Twelve (6 pairs) male
cadaveric lower extremity limbs were externally rotated 30º using two selected football
shoe designs, i.e. a flexible shoe and a rigid shoe. Motion capture was performed to track
the movement of the talus with a reflective marker array screwed into the bone. A
computational ankle model was utilized to input talus motions for the estimation of ankle
ligament strains. At 30º of rotation the rigid shoe generated higher ankle joint torque at
46.2 ± 9.3 Nm than the flexible shoe at 35.4 ± 5.7 Nm. While talus rotation was greater in
the rigid shoe (15.9 ± 1.6º vs. 12.1 ± 1.0º), the flexible shoe generated more talus
eversion (5.6 ± 1.5º vs. 1.2 ± 0.8º). While these talus motions resulted in the same level
of anterior deltoid ligament strain (approx. 5%) between shoes, there was a significant
increase of anterior tibiofibular ligament strain (4.5 ± 0.4% vs. 2.3 ± 0.3%) for the
flexible versus more rigid shoe design. The flexible shoe may provide less restraint to the
subtalar and transverse tarsal joints, resulting in more eversion but less axial rotation of
the talus during foot/shoe rotation. The increase of strain in the anterior tibiofibular
ligament may have been largely due to the increased level of talus eversion documented

110

for the flexible shoe. There may be a direct correlation of ankle joint torque with axial
talus rotation, and an inverse relationship between torque and talus eversion. The study
may provide some insight into relationships between shoe design and ankle ligament
strain patterns. In future studies these data may be useful in characterizing shoe design
parameters and balancing potential ankle injury risks with player performance.

Keywords: biomechanical study; ankle injury; talus eversion; ligament strain;
computational model; motion analysis; ankle kinematics; footwear.

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INTRODUCTION
Acute injuries that occur to the ankle are among the most frequent musculoskeletal
injuries in all levels of sports, and ligament sprains account for 75% of these injuries [1].
In young athletes acute ankle trauma is responsible for 10% to 30% of all sports-related
injuries [2]. Each year an estimated one million persons present to physicians with acute
ankle injuries [3]. Approximately 85% of ankle sprains involve the lateral ankle
ligaments that are ascribed to excessive foot inversion [4, 5]. In contrast, high ankle and
medial ankle sprains occur less frequently, being diagnosed in 10% to 15% of cases [6, 7].
As opposed to a lateral ankle sprain, high and medial ankle sprains are more problematic
due to their potential for resulting in significantly greater time lost and subsequent
chronic ankle dysfunction [7-11]. The mechanism of injury in high and medial ankle
sprains is commonly ascribed to excessive internal rotation of the upper body, while the
foot is planted on the playing surface [8].

Numerous studies have investigated the role of shoe design in the characteristics of shoesurface interfaces [12-16]. While linear traction between a shoe’s outsole and a sports
surface is necessary for high-level performance during any athletic contest, it is generally
accepted that excessive rotational traction (torque) may result in ankle and knee injuries
[15, 17, 18]. Additionally, Livesay et al. [19] also measured the rotational stiffness of
shoe-surface combinations and showed that differences in rotational stiffness are often
greater than differences in peak torque generated between various combinations. The
study concludes that rotational stiffness of the shoe-surface interface may be another
important risk factor for lower extremity injuries. A recent study by Villwock et al. [15],

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involving football shoes and various natural and synthetic playing surfaces, suggests that
the shoe-surface rotational stiffness may be associated, in part, with the design of a shoe’s
upper. However, the effect of shoe design on the patterns of ankle ligament strain during
external rotation of the foot has not been directly, to date, investigated.

Talus motion plays an important role in developing ankle ligament strains, especially
under rotational loading [20], and therefore its motion is crucial in the study of potential
mechanisms of ankle ligament sprain [21]. Recently, Wei et al. [22] developed a
computational ankle model, based on a generic computed tomography scan of a cadaver
foot, which was validated against experimental data from human cadaver ankles [23].
The model has been used to estimate ankle ligament strains and ankle joint torque during
simulations of external foot rotation. In a more recent in vivo study by the same group
[24], barefoot subjects performed single-legged, internal rotation of the body with a
planted foot while a motion capture system tracked motion of the ankle. The kinematic
data of the talus were then utilized to drive the computational model for estimation of the
dynamic ankle ligament strains. The study showed the largest strains in the anterior
tibiofibular (ATiFL) and anterior deltoid (ADL) ligaments, with strain peaking in the
ATiFL prior to the ADL.

The purposes of the current study were (1) to evaluate a few football shoe designs in
order to select two designs of significantly different rotational stiffness (i.e. the highest
and the lowest), then (2) to conduct external rotation tests using human cadaver ankles in
these two shoe designs in order to measure differences in talus motion for the same level

113

of external shoe rotation, and finally (3) to input the talus kinematics into the validated
computational model in order to estimate the differences in the patterns of key ankle
ligament strains between these two extremes in shoe design. It was hypothesized that
shoes with significantly different rotational stiffness would generate significantly
different levels of torque and key ankle ligament strain patterns. These data would begin
to show how differences in shoe design may influence differences in ankle ligament
strains and potentially help guide studies on the role of shoe design in determining some
of the mechanisms of ankle ligament injury during external rotation of the foot.

METHODS
Shoe Stiffness Tests
Four football shoe designs were evaluated to determine their rotational stiffnesses (the
rate at which torque is developed under rotation). Experiments were conducted on a
custom, hydraulic, biaxial testing machine using a 244 Nm rotary actuator (Model SS001-1V, Micromatic, Berne IN) and a vertically oriented linear actuator (Model 204.52,
MTS Corp.). The four football shoe types were Nike Air, Nike Merciless, Adidas Blitz,
and Nike Flyposite. A surrogate lower extremity was made of room temperature curing
epoxy resin (Fiber Strand, Martin Senior Corp., Cleveland, OH) and a stainless steel rod
(Figure 6-1a). The surrogate foot was fitted in a left, size-10 shoe of each design. A
football cleat mold (Figure 6-1b) was made of the same epoxy material for each shoe and
was inserted into an aluminum tray (Figure 6-1c) which was secured to the rotationlocked linear actuator of the test machine with a custom fixture that allowed x-y
adjustments to align the rod along with the linear and torsional actuators. The proximal

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end of the surrogate limb was inserted into an aluminum box (Figure 6-1c) which was
attached to the rotary actuator through a biaxial load cell (Model 1216CEW-2K, Interface,
Scottsdale, AZ) with a capacity of 8896 N axial force and 113 Nm torsion (Figure 6-1c).
A pilot study, using a different load cell and shoes, showed that this football cleat-mold
structure can bear up to 200 Nm torque without observable damages to the shoe or the
mold (data not shown).

Figure 6-1 Shoe stiffness tests preparation and setup. Surrogate lower extremity (a) and
football cleat mold (b) were made of epoxy resin. A surrogate limb was attached to the
testing machine through a biaxial load cell (c).

A compressive pre-load of 1500 N and a rotational pre-torque of 2 Nm were applied to
the surrogate limb prior to internal rotation of the rod (external rotation of the foot). The
magnitude of the pre-load was approximately 2 times body weight and selected to
simulate weight bearing in a dynamic situation [23]. The pre-torque was to ensure full
contact between the medial edge of the foot and the shoe. A dynamic torque of 60 Nm
was input in load control at a frequency of 1 Hz (0.5 s to peak torque) and repeated two

115

more times for each shoe design. The loading portions of the torque-rotation curves were
averaged across the three cycles and compared between shoes. The shoe rotational
stiffness, defined as the slope of the torque-rotation curve (loading portion), was
calculated in Nm/deg and also averaged and compared between shoes.

Cadaver Tests
Twelve (six pairs) fresh-frozen lower limbs from male cadavers (aged 56 ± 12 years)
were used in these tests. The limbs were stored at -20°C and thawed to room temperature
for 24 hours prior to tests. The tibia and fibula were transected approximately 15 cm
distal to the center of the knee. The proximal end of the tibia and fibula shafts were then
cleaned with 70% alcohol and potted in an aluminum box with epoxy resin (Figure 6-2a).
Two screws were placed in the medial and lateral aspects of the proximal tibia, with an
approximately 30 mm projected length, to help prevent the tibia from rotating within the
potting material (Figure 6-2b). From the shoe stiffness tests, shoes with the highest and
lowest rotational stiffnesses were referred as the rigid and flexible shoes, respectively,
and were randomly assigned to the left or right limbs. For each pair of feet one foot was
in the rigid shoe and the other was in the flexible shoe. Shoes were properly selected to fit
the cadaver foot size and regular sports socks were used. The limb was then mounted
upside down (foot pointing upward) in the test machine (Figure 6-2c) with the same
fixture as in the shoe stiffness tests.

The same pre-load (1500 N) and pre-torque (2 Nm) were applied along the axis of the
tibia. Internal tibial rotations (external foot rotations) of 30 degrees were input in position

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control at a frequency of 1 Hz (0.5 s to peak rotation). This rotation magnitude was
selected based on a previous study [23] to ensure that no ankle failure/injury would occur
during these tests. Only one trial for each cadaver foot was performed to eliminate any
viscoelastic and micro-damage effects in the soft tissue. Maximum torques in the 30
degrees of rotation tests were documented for each specimen. The loading portions of the
torque-rotation curves and the slopes of their linear ranges (rotational stiffness in Nm/deg)
were averaged across specimens and compared between shoes. Linear regression was
performed for the linear portion of each curve and the intercept along the rotation axis
was documented and compared between shoes.

Figure 6-2 Cadaver tests preparation and setup. The proximal end of the shank was
potted using epoxy resin (a) with two screws placed earlier into the proximal tibia (b). A
cadaveric limb with markers was mounted upside down into the testing machine (c).

Motion Analysis
The cadaver tests were performed with two reflective marker arrays screwed into the
talus and the tibia, respectively (Figure 6-2c). These arrays were utilized in subsequent
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motion analyses of the talus with respect to the tibia using a five-camera Vicon motion
capture system (Oxford Metrics Ltd., Oxford, United Kingdom) (Figure 6-3). The talus
array was attached from the anterior aspect of the ankle through a hole with a diameter of
approximately 20 mm in the tongue of the shoe. Care was taken to avoid damage to the
anterior deltoid ligament complex (a combination of the anterior tibiotalar and
tibionavicular ligaments). The tibia array was positioned 10-20 cm proximal to its
inferior articular surface (Figure 6-2c). A joint coordinate system (JCS) was established
based on each reflective marker array, as described in previous studies [22, 25, 26]. The
translations and rotations of the talus relative to the tibia in three directions were
determined in this JCS for the 30 degrees of foot rotation test on each limb. Temporal
profiles of talus rotation in different shoes were generated and compared with the actual
shoe rotation.

Model Simulation
Talus motion data were used to drive a generic computational ankle model. Details of
model development and motion simulation have been described in previous studies [22,
24], thus only a brief description is given here. The ankle model was constructed from a
generic computed tomography (CT) scan of a cadaveric ankle with a separation of 0.6
mm between slices. CT images were first converted into 3-D models in MIMICS
(Materialise, Ann Arbor, MI) and then imported into dynamic rigid-body motion
simulation software (SolidWorks, TriMech Solutions, LLC, Columbia, MD). This ankle
model includes 21 ligaments formulated as linear elastic springs with properties adapted

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from the literature [22]. Ligament strains, defined in percentage as the relative
elongations of ligaments, were estimated from the computational model.

5X Vicon
cameras

Figure 6-3 Testing setup. Five-camera Vicon motion capture system (showing only 4
cameras) and one video camera (not shown) were used to track motions of the talus
relative to the tibia.

Statistical Analysis
One-way ANOVA and Student-Newman-Keuls (SNK) post hoc tests were used to
statistically compare differences in rotational stiffness between shoe designs, ankle joint
torques at 30º rotation, shoe/ankle rotational stiffness, intercept along the rotation axis,

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and talus motion (external rotation and eversion) relative to the tibia between the paired
limbs. Two-way ANOVA and SNK post hoc tests were used to determine the differences
in the torque-rotation data between shoe designs (factor one) at each torque level (factor
two) for the shoe stiffness tests, between limbs (factor one) at each rotation level (factor
two) for the cadaver tests, and the difference in ligament strains between the flexible and
rigid shoes (factor one) in different ligaments (factor two) for the model simulation. In all
statistical tests p values less than 0.05 were considered significant.

RESULTS
Shoe Stiffness Tests

Figure 6-4 Torque-rotation curves of the four shoe designs from rotational stiffness tests.
Different symbols (‡ # § *) indicate statistically significant differences between shoe
designs.
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The pre-torque of 2 Nm was zeroed out by shifting all curves downward in both Figure 64 for the shoe tests and Figure 6-6 for the cadaver tests. Torque-rotation curves
demonstrated nearly linear behavior and significant differences between the four shoe
designs (Figure 6-4). While the Air design showed the lowest rotational stiffness (21.9 ±
2.8 Nm/deg), the Flyposite design had the highest stiffness (50.0 ± 1.7 Nm/deg) (Figure
6-5). These two kinds of shoes were then used in the cadaver tests with the Air design
being referred as the flexible shoe and the Flyposite design as the rigid shoe. The
Merciless and Blitz designs also showed a significant difference in rotational stiffness
(34.5 ± 1.9 Nm/deg versus 37.4 ± 2.1 Nm/deg, respectively).

Figure 6-5 Rotational stiffness determined from the slopes of torque-rotation curves.
Data were averaged across the three cyclic tests and plotted as mean ± 1 SD. The Air was
the least rigid (most flexible) shoe, while the Flyposite was the most rigid design.
Different symbols (‡ # § *) indicate statistically significant differences between shoe
designs.
121

Cadaver Tests

Figure 6-6 Torque-rotation curves of cadaveric limbs restrained by the flexible (Air) or
the rigid (Flyposite) shoe designs showed a toe region followed by a linear region. Linear
regression was performed on the linear portion of the curves between 10° and 30°. §
indicates significant difference between curves. ‡’s indicate significant differences
between limbs at various rotation points.

There was an obvious toe region in the torque-rotation curves for both shoe tests (Figure
6-6). No significant changes in torque values were observed during the first 5° of external
rotation. Torque-rotation responses overall were significantly different between limbs,
with the limb in the rigid shoe stiffer than that in the flexible shoe (Figure 6-6). The
122

shoe/ankle rotational stiffness, defined as the slope of the linear portion of the torquerotation curves between 10º and 30º (Figure 6-6), was significantly greater for the rigid
shoe (1.96 ± 0.24 Nm/deg) than for the flexible shoe (1.65 ± 0.18 Nm/deg) (Table 6-1).
Intercept of linear regression along the rotation axis (Figure 6-6) was statistically
different between the flexible (8.8 ± 1.2º) and rigid shoes (6.2 ± 0.9º) (Table 6-1). At 30º
of external foot rotation, ankle joint torque in the rigid shoe (46.2 ± 9.3 Nm) was
statistically higher than that in the flexible shoe (35.4 ± 5.7 Nm) (Table 6-1).
Table 6-1 Specimen descriptions and results from the cadaver tests.

Spe# Age

Height
(m)

Weight
(kg)

Torque at 30°
rotation (Nm)a
Flexible

Rigid
58.8
(R)
42.4
(L)
40.2
(R)
42.8
(L)
36.4
(R)
56.8
(L)

Rotational
stiffness
(Nm/deg)b

Intercept along x
axis (deg)c

Flexible

Rigid

Flexible

Rigid

1.75

2.13

9.5

6.6

1.48

1.68

8.9

6.4

1.71

2.18

7.2

5.3

1.84

2.02

10.6

7.7

1.39

1.64

7.7

5.2

1.74

2.09

8.6

5.9

1

76

1.72

64

39.5 (L)

2

56

1.83

79

36.9 (R)

3

40

1.88

75

34.5 (L)

4

50

1.75

88

37.9 (R)

5

55

1.78

66

24.3 (L)

6

56

1.88

93

39.2 (R)

56

1.81

77.5

35.4

46.2*

1.65

1.96*

8.8

6.2*

12

0.07

11.6

5.7

9.3

0.18

0.24

1.2

0.9

Mea
n
SD
Note:

a

The left (L) and right (R) limbs were randomly assigned.
Rotational stiffness was defined as the slope of the linear portion of the torquerotation curves (between 10º and 30º) in Figure 6-6.
c
Linear regression of each curve in Figure 6-6 was intercepted with the rotation
(x) axis and the intercept was documented.
* Statistically different than in the flexible shoe (p<0.001).
b

123

Motion Analysis
During axial loading of the ankle and prior to external rotation of the foot, talus eversion
was noted in both shoes without a significant difference (1.4 ± 0.5º for the flexible shoe
versus 1.3 ± 0.5º for the rigid shoe) (Table 6-2). In the 30º of external foot rotation test,
the talus externally rotated more in the rigid shoe (15.9 ± 1.6º) than in the flexible shoe
(12.1 ± 1.0º). Talus eversion, however, was found to be significantly greater in the
flexible shoe (5.6 ± 1.5º) than in the rigid shoe (1.2 ± 0.8º) (Table 6-2). Medial-lateral
translation of the talus was minimal in both shoes (< 1 mm). In addition, external rotation
of the talus during axial loading and talus dorsi/plantar flexion were also small and
negligible (< 0.5 deg). Temporal profiles showed a statistically different talus rotation in
different shoes (p < 0.001), with both talus rotations much less than the shoe rotation
(Figure 6-7).
Table 6-2 Talus motion relative to the tibia (mean ± SD) in different shoes.

Talus motion a

Axial loading of 1500 N

Foot at 30° external rotation

Flexible
External rotation (deg)
Eversion (deg)
Note:

Rigid

Flexible

Rigid

---

---

12.1 ± 1.0

15.9 ± 1.6*

1.4 ± 0.5

1.3 ± 0.5

5.6 ± 1.5

1.2 ± 0.8*

a

Talus external rotation during axial loading (---) and talus dorsi/plantar flexion
(not shown) were minimal and negligible (< 0.5°).
* Statistically different than in the flexible shoe (p<0.001).

Model Simulation
While ligament strains in the ADL were at the same level for both shoes (4.9 ± 0.4% for
the flexible shoe versus 5.2 ± 0.7% for the rigid shoe), the flexible shoe generated
significantly higher strains in the ATiFL (4.5 ± 0.4%) than the rigid shoe (2.3 ± 0.3%)
124

(Figure 6-8). For both shoes, the ADL experienced higher strains than the ATiFL
(p=0.043 for the flexible shoe; p < 0.001 for the rigid shoe).

Figure 6-7 Comparisons of temporal profiles of external talus rotations in different shoes
with the actual shoe rotation (same in all tests) driven by the rotary actuator. All rotations
were relative to the tibia.

125

Figure 6-8 Ankle ligament strains (mean ± 1 SD) were estimated from a computational
model and compared between different shoes at 30° of external foot rotation. Only two
ligaments with the highest strains were reported. ATiFL is the anterior tibiofibular
ligament, and ADL is the anterior deltoid ligament. The horizontal bar indicates
significant difference between shoes. The strain in the ADL was statistically greater than
in the ATiFL for both shoe designs.

DISCUSSION
This study investigated the rotational stiffness of four football shoe designs. The most
flexible and rigid shoes were then used in cadaver experiments to investigate the effects
of this shoe property on ankle joint torque and talus motion during external rotation of the
foot. The talus kinematic data were then input into a generic, computational ankle model
for determination of key ankle ligament strains. The results supported the hypotheses that
shoes with different rotational stiffness would generate different levels of ankle joint

126

torque and ankle ligament strain patterns. For a given level of external foot rotation the
rigid shoe developed more torque than the flexible shoe. Furthermore, there was a
dramatic jump in strain of the ATiFL for the flexible versus the rigid shoe design, while
the ADL strain was nearly the same in both shoes. The significant increase in ATiFL
strain may be largely due to the greater talus eversion documented for the flexible shoe.
These results are supported, in part, by an earlier study from our laboratory showing that
during external rotation tests a shoe with a more pliable upper tended to have its medial
edge dig into the turf more than a rigid shoe design [15]. This motion of the more flexible
shoe may allow the foot to evert more than a rigid design during external rotation of the
foot. In another more recent study by this laboratory eversion of an axially loaded foot
prior to external rotation transferred the site of ligament failure from the ADL to the
ATiFL [27]. While all specimens in the current study were initially placed in a neutral
position, during external rotation of the foot there was significantly more coupled
eversion of the talus for the flexible shoe design. As that might have been expected with a
neutral foot, both shoe designs showed the largest strains in the ADL.

The mean ankle ligament strains in the ATiFL and ADL from the model analysis of the
current study indicated levels of approximately 2.3% and 5.2%, respectively, for the rigid
shoe under 30º of foot rotation. Interestingly, a previous simulation study in which the
cadaver foot was restrained with athletic tape showed a comparable pattern of ligament
strains with the ATiFL at 2% and the ADL at 7% under the same conditions of axial load
and foot rotation as for the rigid shoe in the current study [22]. A previous study by
Verhagen et al. [28], investigating the efficacy for preventive measures on ankle sprains,

127

shows that while the use of tape reduces the incidence of ankle sprains, the efficacy of
shoes in preventing the injury is unclear. The data from the current study, however, may
suggest some level of parallel restraint to the foot between this method of ankle taping
and the more rigid football shoe design.

In the current study, while the flexible shoe generated less external rotation of the talus
relative to the tibia, it resulted in significantly greater talus eversion than using the rigid
shoe design. One potential explanation for these contrasting talus motions may be that the
more flexible shoe provided less restraint to the subtalar and transverse tarsal joints than
the more rigid shoe. This allowed relatively more motion between the talus and the
calcaneus, generating foot eversion [21] for the flexible shoe. And, the flexible shoe also
allowed more motion between the forefoot and the hindfoot, producing less axial talus
rotation during external rotation of the foot. In contrast, the more rigid shoe design
coupled motion of the foot with the talus during external rotation, but might prevent the
coupling of talus eversion with external rotation of the foot that would be expected with
an unrestrained foot and ankle condition [29]. Unfortunately, motions of the calcaneus,
navicular and cuboid in the shoes were not measured in the current study due to the
difficulty in placing marker arrays into those bones. As a result subtalar and transverse
tarsal joint motions were unknown. Cutting holes in the shoes may be a solution in future
studies, assuming the consequent alteration in shoe property was minimal.

The current study also suggested a correlation between ankle joint torque and axial talus
rotation. For example, at 30º of foot/shoe rotation, axial talus rotation in the flexible shoe

128

was 12.1°, generating 35.4 Nm of ankle joint torque. In contrast, axial talus rotation in
the rigid shoe was 15.9°, generating 46.2 Nm of ankle joint torque. Interestingly, the ratio
of ankle joint torque to axial talus rotation was similar at 2.9 Nm/deg between shoe
designs. This stiffness, which seemed to be independent of shoes, but likely dependent on
the given conditions of ankle joint pre-load and rate of external foot rotation, may
characterize a structural property of the human ankle. An earlier study by this laboratory
shows that the in vivo foot rotational stiffness is 1.1 – 1.5 Nm/deg in terms of axial
hindfoot rotation, which might be expected, however, to be greater than talus rotation at
the same level of foot rotation [24].

A previous study by Reinschmidt et al. [30] compares tibiocalcaneal motion during
running based on skeletal markers with the motion based on external markers attached to
the shoe and shank. The study shows that the mean difference between external and
skeletal marker-based rotations is 51.2% of the total motion, and therefore concludes that
rotations derived from external shoe and shank markers typically overestimate the
skeletal tibiocalcaneal kinematics. In contrast, the current study was designed to measure
tibiotalar motion during external foot rotation. While no external markers were attached
to the shoes in the current study and shoe motion was restrained to internal-external
rotation only, shoe rotation was accurately controlled by the testing fixture. Interestingly,
the current study also demonstrated a large difference between shoe rotation and talus
rotation. For example, at 30° of rigid shoe rotation the average talus rotation was
approximately 15.9°. In contrast, with the more flexible shoe restraint 30º of shoe rotation
generated approximately 12.1º of talus rotation, showing that shoe design can yield

129

differing level of restraint to the ankle and may play a role in the potential for ankle joint
injury.

While the current studies were limited in that only a few shoe designs were considered
and the study was of a subfailure nature, a couple future studies could be envisioned.
First, the current study showed that external rotation of the foot/shoe to 30º generated
ankle ligament strains up to approximately 5%. Future cadaver studies, possibly with a
similar experimental setup to vary the levels of foot dorsiflexion and eversion, could
rotate the ankle to a failure level to document some effects of football shoe design on the
potential for and subsequent location of ankle injury. Secondly, the current experimental
setup could be modified to study shoe-surface interface characteristics in order to
investigate the potential influence of shoe-surface interface in developing ankle ligament
strains as an indication of the potential for location and severity of ankle ligament injury,
similar to that previously conducted to study the role of shoe-surface interface on the
potential for knee ligament injury [31].

In conclusion, we externally rotated six pairs of cadaver limbs in two different football
shoe designs, a flexible shoe and a rigid shoe, and found that while axial talus rotation
was significantly greater in the rigid shoe, the flexible shoe generated statistically more
talus eversion. The study showed that football shoe design can have an effect on the
pattern of ankle ligament strains during external rotation of the foot to potentially
influence the location and severity of a subsequent ankle injury. While the current study
was indeed limited in scope, it represents a first step in our attempt to understand the

130

effect of football shoe design on the potential for ankle injury. These studies must also
consider the possible implications of football shoe design on player performance and
balance this factor with the potential for ankle injury.

131

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132

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135

CHAPTER 7
A COMPUTATIONAL MODEL TO INVESTIGATE SHOE AND SHOE-SURFACE
INTERFACE EFFECTS ON ANKLE LIGAMENT STRAINS DURING A
SIMULATED SIDESTEP CUTTING TASK

ABSTRACT:
Ankle sprains account for 10% to 15% of reported sports injuries. High ankle sprains are
currently thought due to torsional loads and potentially debilitating to the athlete. In the
current study a computational model was developed to investigate the human response in
shoes on different athletic playing surfaces during a simulated sidestep cutting task.
Ankle ligament strains were obtained from the model to help predict ankle injury. The
model may provide a computational basis for studying shoes and shoe-surface interfaces
that can be used to help optimize player performance and minimize injury risk.

KEY WORDS: ligament, ankle injury, foot constraint, shoe-surface interface, model.

136

INTRODUCTION:
Ankle sprains are one of the most frequent injuries in sports and often account for 10% to
15% of reported injuries [7]. In the National Football League high ankle sprains, also
known as syndesmotic sprains, constitute approximately 20% of ankle injuries [2]. These
injuries receive considerable attention due to their relatively long recovery time. While
high ankle torque is implicated as a risk factor, a cadaver study by Villwock et al. [7]
suggests foot rotation may be a better predictor of injury. In addition, Villwock et al. [8]
investigated the effects of various cleated shoe designs and playing surfaces on torsional
responses using a surrogate ankle. Their study shows that synthetic surfaces yield higher
torques and rotational stiffnesses than natural grass surfaces. The study also suggests that
a shoe with a pliant upper, allowing more subtalar motion, may provide less foot
constraint than a rigid upper, and consequently help protect the foot from rotational
injuries.

The objective of current study was to develop a computational model, guided by the
above cadaver and playing surface studies, to investigate the effects of foot constraint and
shoe-surface interface on ankle ligament strains. The model may have utility in providing
a computational basis for shoe and playing surface designs that will optimize player
performance and minimize injury risk.

METHOD:
Joint anatomical features were taken from one computed tomography (CT) set in the
cadaver studies [7]. Detailed features of the ankle were obtained by importing Digital

137

Imaging and Communications in Medicine (DICOM) files of individual CT scans into
Materialise’s Interactive Medical Imaging Control System (MIMICS) (Materialise, Ann
Arbor, MI). This yielded a three-dimensional surface model of the bones as
Stereolithography (STL) files. To reduce the size of the surface files and subsequent
model, the STL files were remeshed in MIMICS to smooth the surfaces of each bone.
Exported files were then imported into the 3-D solid modelling software SolidWorks
(TriMech Solutions, LLC, Columbia, MD) as Mesh Files (.stl) [5]. SolidWorks, along
with its ScanTo3D package, was used to further construct each bone and simplify the
bone surfaces. A rigid plane (200 x 100 x 10 mm) was created in SolidWorks to represent
the playing surface. The SolidWorks Motion package was then used to assemble the
bones and surface, obtain proper positioning, add necessary components, and run
simulations.

The ligaments were represented as linear springs (Fig 7-1), and their stiffness values,
origins and insertions were based on the literature [5,6]. The foot model was composed of
14 bones. A compressive pre-load of 1600 N, approximately 2X BW, was distributed
between the tibia and fibula. Two kinds of foot constraint were simulated, representing
rigid and pliant shoes. The rigid shoe model had 3 contacts and 10 ligaments. The tibia
was fixed in space, and the fibula and remaining bones (acting as a rigid body) were free
to move (Fig 7-1 left). The pliant shoe model had 6 contacts and 12 ligaments. The tibia
was fixed in space, and the fibula, calcaneus and the remaining bones (acting as a rigid
body) were free to move, allowing subtalar motion (Fig 7-1 right).

138

Figure 7-1: Foot and surface models. Left: rigid shoe; Right: pliant shoe.

The shoe-surface interface was simulated with a torsional spring at the surface center and
3 lateral surface springs, as shown in Figure 7-1 (red springs). Two kinds of surfaces,
synthetic turf and natural grass, were simulated with stiffnesses of 3.6 Nm/deg and 2.2
Nm/deg, respectively, based on the literature [8]. The stiffness of the 3 lateral surface
springs was 1 N/mm. A rotary motor was applied to the surface aligned at the ankle
center between the malleoli. The external rotation angle was set to 30° to simulate a
common sidestep cutting task [3].

RESULTS: Critical ligament strains were measured in the computational models for
various test conditions. For clarity, only strains of the ligaments injured in the previous
cadaver experiments [7] were plotted in Figure 7-2 and Figure 7-3; namely, the posterior
talofibular ligament (PTaFL), the deltoid ligament (representing the anterior tibiotalar
ligament), and the anterior tibiofibular ligament (ATiFL).

139

The maximum strains in each ligament were greater on the synthetic turf than on natural
grass, and the order, from the highest to the lowest, was rigid shoe on turf, pliant shoe on
turf, rigid shoe on grass, and pliant shoe on grass (Fig 7-2). This indicated that shoesurface interface influenced ankle ligament strain and consequently the potential for
ankle injury. For the rigid shoe condition, the maximum strain occurred in the PTaFL,
while the deltoid ligament experienced the highest strain in the pliant shoe, suggesting
that the location of ankle injury might depend on foot constraint. On the same surface,
higher ligament strains were developed for the rigid than pliant shoe, indicating that less
rotation might be required to fail ligaments with a rigid upper design. Finally, two large
changes in ligament strain were noted in the PTaFL, namely, from Rigid/Turf to
Pliant/Turf and from Rigid/Grass to Pliant/Grass. Apparently, these were due to foot
constraint, and therefore, the effect of foot constraint on ligament strain seemed to be
greater for the PTaFL than the other two ligaments. Comparably, significant changes due
to surface were noted in both the PTaFL and the deltoid ligament, such as from
Rigid/Turf to Rigid/Grass and from Pliant/Turf to Pliant/Grass. This finding indicated
that shoe-surface interface affected strains in the PTaFL and deltoid ligament in a similar
fashion.

The strain-rotation behaviours of each ligament were plotted in Figure 7-3. While linear
springs were used to represent ligaments in these models, the strain-rotation behaviour
was non-linear due to the complex joint geometry.

140

PTaFL

Deltoid

ATiFL

Max Ligament Strain (%)

35
R/T---Rigid/Turf
P/T---Pliant/Turf
R/G---Rigid/Grass
P/G---Pliant/Grass

30
25
20
15
10
5
0
R/T

P/T

R/G

P/G

R/T

P/T

R/G

P/G

R/T

P/T

R/G

P/G

Shoe-Surface Interface

Figure 7-2: Strains in various ligaments at 30° of external surface rotation for different

Ligament Strain (%)

shoes and shoe-surface interface conditions.

40

Rigid / Turf

40

Pliant / Turf

40

Rigid / Grass

40

30

30

30

30

20

20

20

20

10

10

10

Pliant / Grass

10

0
0

10

20

30

0
40 0

External Rotation
(deg)

10

20

30

0
40 0

External Rotation
(deg)

10

20

30

PTaFL
Deltoid
ATiFL

0
40 0

External Rotation
(deg)

10

20

30

40

External Rotation
(deg)

Figure 7-3: Strains in various ligaments during external rotation of the surface for
different shoes and shoe-surface interface conditions.

DISCUSSION:
The results from the current models were consistent with ligament injuries in the cadaver
study [7], suggesting that ankle injury depends on foot constraint. Models of the shoe and
141

shoe-surface interface conditions in the current study indicated that larger strains were
developed in ankle ligaments on synthetic turf than on natural grass for a simulated
sidestep cutting task.

Subtalar motion of the foot was allowed in the pliant shoe model by freeing up the
calcaneus (Fig 7-1 right). Consequently, ankle ligaments in the pliant shoe experienced
less strain than in the rigid shoe on the same surface. This agreed with the Villwock et al.
[8] study suggesting that a pliant shoe upper may help reduce the risk of ankle injury. For
performance purposes, however, football and soccer players tend to wear tight fitting
shoes with relatively rigid uppers to provide more foot constraint. An in vivo study by
Dowling et al. [3] showed that during the sidestep cutting task, subjects were able to
obtain the desired cut of approximately 30° on the high friction surface but only 24° on
the low friction surface. Their study, together with Villwock et al. [8], suggests that
synthetic surfaces with high friction may benefit player performance. Yet, according to
the current study, the synthetic turf may compromise ankle mechanics and increase the
risk of injury.

A recent study by Drakos et al. [4] investigated the effect of different shoe-surface
combinations on ACL strain and suggests that a cleat-grass interface may result in fewer
noncontact ACL injuries than the turf shoe-turf interface. Similarly, Dowling et al. [3]
suggests high friction synthetic turfs may be associated with an increased incidence of
ACL injury. While the current study documented only ankle ligament strains, the results

142

support the notion that synthetic surfaces may be a potential risk factor for rotational
injuries of the lower extremity, especially when using a rigid shoe design.

Some limitations of the model should be noted. A study by Beumer et al. [1] determined
the strength and stiffness of the tibiofibular and tibiotalar ligaments of the ankle and
showed no differences between these ligaments having an average strength and stiffness
of 550 N and 98 N/mm, respectively. Ligament elongation at failure was estimated to be
approximately 5.6 mm. Based on an estimated length of these ligaments from the current
study, this level of elongation may suggest failure strains on the order of 25% to 31%.
While this level of strain was predicted by the rigid shoe and synthetic turf condition for
the simulated sidestep cutting task, our model did not include any nonlinearity in
response of the ligaments. We would suggest future simulations incorporate this well
documented nonlinear response of ankle ligaments for failure prediction studies. While
cleat pattern at the shoe-surface interface is an important factor in shoe design, its effect
on failure analyses is yet unknown and should be studied in the future.

CONCLUSION:
Computational models of shoe and shoe-surface interface conditions were developed to
measure ankle ligament strains during a simulated sidestep cutting task. The results
showed that the maximum ankle ligament strains were generated with a rigid shoe model
on synthetic turf. The ability of the current models to incorporate various shoe-surface
interface characteristics and show differences in predicted ankle ligament strains was
encouraging. While more failure data are needed on ankle ligaments, additional

143

experiments using human cadavers and in vivo tests with human subjects will also be
needed to validate computation models. Ultimately, such models may provide a basis for
optimizing shoe designs and shoe-surface interface characteristics to enhance player
performance and minimize injury risk.

144

REFERENCES

145

REFERENCES

1. Beumer, A., van Hemert, W.L., Swierstra, B.A., Jasper, L.E., & Belkoff, S.M.
(2003). A biomechanical evaluation of the tibiofibular and tibiotalar ligaments of
the ankle. Foot Ankle Int., 24(5), 426-429.
2. Boytim, M.J., Fischer, D.A., & Neumann, L. (1991). Syndesmotic ankle sprains.
The American Journal of Sports Medicine, 19(3), 294-298.
3. Dowling, A.V., Corazza, S., Chaudhari, A.M.W., & Andriacchi, T.P. (2010).
Shoe-surface friction influences movement strategies during a sidestep cutting
task: implications for Anterior Cruciate Ligament injury risk. The American
Journal of Sports Medicine, 38(3), 478-485.
4. Drakos, M.C., Hillstrom, H., Voos, J.E., Miller, A.N., Kraszewski, A.P.,
Wickiewicz, T.L., Warren, R.F., Allen, A.A., & O’Brien, S.J. (2010). The effect
of the shoe-surface interface in the development of Anterior Cruciate Ligament
strain. ASME Journal of Biomechanical Engineering, 132, 011003.
5. Liacouras, P.C., & Wayne, J.S. (2007). Computational modelling to predict
mechanical function of joints: application to the lower leg with simulation of two
cadaver studies. ASME Journal of Biomechanical Engineering, 129, 811-817.
6. Netter, F.H., & Hansen, J.T. (2003). Atlas of Human Anatomy. 3rd ed. Icon
Learning Systems, Teterboro, NJ.
7. Villwock, M.R., Meyer, E.G., Powell, J.W., & Haut, R.C. (2009a). External
rotation ankle injuries: investigating ligamentous rupture. ASME Summer
Bioengineering Conference, Jun 17-21, 2009, Lake Tahoe, CA.
8. Villwock, M.R., Meyer, E.G., Powell, J.W., Fouty, A.J., & Haut, R.C. (2009b).
Football playing surface and shoe design affect rotational traction. The American
Journal of Sports Medicine, 37(3), 518-525.

146

CHAPTER 8
DETERMINATION OF DYNAMIC ANKLE LIGAMENT STRAINS FROM A
COMPUTATIONAL MODEL DRIVEN BY MOTION ANALYSIS BASED
KINEMATIC DATA

ABSTRACT
External rotation of the foot has been implicated in high ankle sprains. Recent studies by
this laboratory, and others, have suggested that torsional traction characteristics of the
shoe-surface interface may play a role in ankle injury. While ankle injuries most often
involve damage to ligaments due to excessive strains, the studies conducted by this
laboratory and others have largely used surrogate models of the lower extremity to
determine shoe-surface interface characteristics based on torque measures alone. The
objective of this study was to develop a methodology that would integrate a motion
analysis-based kinematic foot model with a computational model of the ankle to
determine dynamic ankle ligament strains during external foot rotation. Six subjects
performed single-legged, internal rotation of the body with a planted foot while a markerbased motion analysis was conducted to track the hindfoot motion relative to the tibia.
These kinematic data were used to drive an established computational ankle model.
Ankle ligament strains, as a function of time, were determined. The anterior tibiofibular
ligament (ATiFL) experienced the highest strain at 9.2 ± 1.1%, followed by the anterior
deltoid ligament (ADL) at 7.8 ± 0.7%, averaged over the six subjects. The peak ATiFL
strain occurred prior to peak strain in the ADL in all cases. This novel methodology may
provide new insights into mechanisms of high ankle sprains and offer a basis for future
evaluations of shoe-surface interface characteristics using human subjects rather than
mechanical surrogate devices.
147

Keywords: ankle sprain, syndesmotic injury, three-dimensional reconstruction, rigidbody model, anterior tibiofibular ligament.

148

INTRODUCTION
Numerous studies have been conducted to evaluate the properties of shoe-surface
interface conditions on natural grass and synthetic turfs [4,21,31,48]. It has been
suggested that ankle injuries correlate with torque generated in the joint. Our laboratory
has developed a surrogate ankle [49] for use with a rotational traction measurement
apparatus to evaluate shoe-surface interface characteristics [48]. External rotation,
commonly thought to be a mechanism of high ankle sprain, is applied to the surrogate
foot and resistive torque is measured. The study suggests that synthetic turfs generate
higher ankle torques than natural grasses, implying a higher injury risk on turfs [5,36].
One limitation of such studies is that artificial devices are used to measure resistive
torques, not strains produced in ankle ligaments. But, high ankle sprains are the result of
injury to the distal tibiofibular ligaments [6,34,38,45,52,53,55].

Marker-based, multi-segmented kinematic foot models have been proposed to evaluate
the motion of bones in the foot during gait [9,25,26,28,42]. The Oxford Foot Model
(OFM), developed by Carson et al. [9], assumes rigid segments including the tibia, the
hindfoot (calcaneus and talus), the forefoot, and the hallux [11,29,44]. Various
computational models have provided considerable insight into the biomechanics of the
ankle joint. Finite element models can provide information about stresses and strains in
bones and ligaments [2,7,8], and multi-body, rigid models have the ability to rapidly
solve for motion-based mechanics of bones within joints [30].

149

The ankle is a common site of acute musculoskeletal injury [53], and sprains account for
75% of these injuries [1]. Approximately 80% of ankle sprains are lateral ankle sprains
caused by excessive inversion [18] while high ankle sprains comprise about 10% [6].
Some studies have examined ankle ligament strains using cadavers [10,39,47]. Using a
combined dual-orthogonal fluoroscopic and magnetic resonance imaging (MRI)
technique, De Asla et al. [12] measure in vivo changes in length of the anterior talofibular
ligament (ATaFL) and calcaneofibular ligament (CaFL) in four static foot positions.
There are limited data, however, on ankle ligaments strains during more sport-specific
dynamic movements. Dynamic ligament strain data may be useful in predicting which
motions predispose ankle soft tissues to injury [46]. The purpose of the current study was
to obtain motion analysis-based data during a single-legged, external foot rotation test,
and use the data to drive a dynamic computational model for estimating ankle ligament
strains. Since high ankle sprain has classically been ascribed to excessive external
rotation of the foot with dorsiflexion and eversion [53,57], the hypothesis of the study
was that the highest ankle ligament strain would be in the anterior tibiofibular ligament
(ATiFL).

METHODS
Institutional Review Board approval was obtained for these experiments. After obtaining
informed consent, kinematic data were collected from six young (age of 27 ± 6 yrs), male
subjects (body weight of 81.2 ± 17 kg) with no history of lower extremity injury. A sixcamera Vicon MX Motion Capture System (OMG plc., Oxford, UK) was used to capture
motion at 100 Hz from a marker set defined by the OFM (OMG plc., Oxford, UK). An

150

AMTI force plate (Advanced Medical Technology, Inc., Watertown, MA) captured
ground reaction data at 1000 Hz. The subjects stood with their right foot planted on the
force plate (Fig. 8-1), and performed an internal rotation of their body (external rotation
of the foot) at their self-determined maximum effort. Each subject was allowed to rotate
at a self-selected speed, but was required to finish the rotation in three seconds. The data
from four trials were processed using commercial kinetic (Plug-in Gait, OMG plc.,
Oxford, UK) and kinematic (Dynamic OFM) packages. A repeatability analysis was
conducted by synchronizing the torque and angle data of the first three trials with the
fourth trial and calculating variance of the four trials for each subject.

Figure 8-1 Marker set of the Oxford Foot Model [9]. The subjects performed a singlelegged, internal rotation of their bodies (arrow) with a planted foot.

A three-dimensional, multibody computational ankle model has been previously
developed by this laboratory [50] based on a generic computed tomography (CT) scan
(0.6 mm per slice) and validated for measurements of ankle ligament strain [10] and
ankle joint torque [51]. Detailed features of the foot/ankle complex were obtained by
importing Digital Imaging and Communications in Medicine (DICOM) files into
151

Materialise’s Interactive Medical Imaging Control System (MIMICS) (Materialise, Ann
Arbor, MI). This yielded a three-dimensional surface model of the bones as
Stereolithography (STL) files for export. The STL files were imported into 3D solid
modeling software (SolidWorks, TriMech Solutions, LLC, Columbia, MD) as Mesh Files.
SolidWorks, along with the ScanTo3D package, constructed each bone and simplified
their surfaces. SolidWorks Motion was used to assemble the bones, obtain proper
positioning, add necessary components (e.g., spring, force, contact, motion, et al.), and
run the simulations. Seventeen ligaments were included and represented as linear spring
elements (Fig. 8-2), with stiffness values from the literature [30,50]. An initial strain of
2% (implemented by inserting a spring element of length 2% shorter than the distance
between insertion points) was assigned to the ankle ligaments based on a previous
cadaver study [37], thereby inducing in situ preloads in the ligaments [54]. An initial
strain of 0.5% was applied to the interosseous ligaments [30]. A model sensitivity
analysis has been conducted in Wei et al. [50] to ensure that moderate ligament stiffness
variations do not significantly alter model conclusions. The tibia was fixed in space, the
fibula was free, and the talus and calcaneus were fused together, as assumed in the OFM
[9]. The remaining bones of the foot were also fused.

Motion data from the OFM were used as input for the simulations. Hindfoot rotations
relative to the tibia as a function of time in three directions (i.e. dorsi-plantar flexion,
inversion-eversion, and internal-external rotation) were determined from the tests. Joint
kinematics from the OFM was deduced by an established joint coordinate system
[19,43,56]. These kinematic data, averaged across the four trials, and body weight of each

152

subject were applied to the computational model. The Akima Spline interpolation method
was used to input continuous rotation-time data into the SolidWorks Motion to drive
hindfoot movement in the model. The body weight was simulated as an axial load and
applied to the proximal end of the model during simulation, distributing one-sixth of the
load on the fibula and the rest on the tibia [27,50]. Resistive moments in the above three
directions were obtained from the simulation and, for purpose of validation, compared
with the moments determined by inverse dynamics from the Plug-in Gait package.
Ligament strains, defined in percentage as the relative elongations of ligaments, were
determined from the computational model, and their dynamic behavior was compared
with ankle motions.

153

a

IOL-I

b

c

IOL-II
ATiFL

ATiTL#
PTiFL

PTiTL*

ATaFL
PTaFL

TiCL*

CaFL

TiNL#
TaNL

CaNL

CaCL
MPF

CPF

LPF

Figure 8-2 Lateral (a), posterior (b), and medial (c) views of a generic ankle model (left limb) showing the locations of 17 modeled
ligaments: the interosseous ligaments (IOL-I and IOL-II); the anterior and posterior tibiofibular (ATiFL and PTiFL); the
calcaneofibular (CaFL); the anterior and posterior talofibular (ATaFL and PTaFL); the anterior and posterior tibiotalar (ATiTL and
PTiTL); the tibionavicular (TiNL); the tibiocalcaneal (TiCL); the talonavicular (TaNL); the calcaneocuboid (CaCL); the
calcaneonavicular (CaNL); the medial, central, and lateral plantar fascia (MPF, CPF, and LPF). The anterior and posterior deltoid
ligaments (ADL and PDL) were each composed of two bundles denoted by # for ADL and * for PDL. Ligament stiffnesses were
documented in Wei et al. [50] with the TiCL stiffness from Liacouras and Wayne [30].

154

Variance of the four trials for each subject was calculated in Excel for repeatability
analysis. One-way ANOVA and Student-Newman-Keuls (SNK) post hoc tests were used
to statistically compare the differences between the maximum strains in ligaments.
Kruskal-Wallis one-way ANOVA on ranks and SNK post hoc tests were used to
determine the differences between temporal sequences of peak ankle motions and peak
ligament strains, with p values less than 0.05 considered significant in both statistical
tests.

RESULTS
The repeatability analysis showed that, for each subject, the variance in hindfoot external
rotation during loading part (Fig. 8-3b) was within 3.2° of the average response curve at
each time point (Table 8-1). And yet, during the unloading phase the variance was
dramatically larger at approximately 16° at each time point.

Table 8-1 Variance (VAR) of the four trials from each subject during hindfoot external
rotation (deg). Rotation profile was separated into loading (initial – peak) and unloading
(peak – end) parts (Fig. 8-3b).
Subjects
Loading VAR
Unloading VAR

1

2

3

4

5

6

Mean

1.9
15.4

3.2
20.1

2.1
15.7

1.5
11.3

1.7
16.2

2.2
14.8

2.1 (deg)
15.6 (deg)

155

Internal (+) / External (-)
Moment (Nm)

a

25
sync point

15
5
-5
trial 1
trial 2
trial 3
trial 4

-15

Time (s)

-25
0

0.5

1

1.5

2

2.5

3

10
sync point

Internal (+) / External (-)
Rotation (deg)

b
0

loading

-10
-20
unloading

trial 1
trial 2
trial 3
trial 4

-30

Time (s)

-40
0

0.5

1

1.5

2

2.5

3

Figure 8-3 Typical profiles of external moment (a) and rotation (b) from one subject.
Data from the first three trials were synchronized with the fourth trial at the peak internal
moment (a).
156

20
10
Time (s)
0
0

1

1.5

2

2.5

average
SD
simulation

20
10
Time (s)
0
0

0.5

1

1.5

2

2.5

0
-5
-10

3

average
SD

-15

Time (s)

-20

10
0
-10
-20
average

-30

SD

Time (s)

-40
0

e

40
30

5

3

Inversion (+) / Eversion (-)
Moment (Nm)

Dorsi (+) / Plantar (-)
Moment (Nm)

d

0.5

Internal (+) / External (-)
Rotation (deg)

average
SD

30

c

10

0.5

1

1.5

2

2.5

3

0

f

15

Internal (+) / External (-)
Moment (Nm)

b

40

Inversion (+) / Eversion (-)
(deg)

Dorsi (+) / Plantar (-)
Flexion (deg)

a

average
SD
simulation

10
5
0
-5

Time (s)
-10

0.5

1

1.5

2

2.5

3

30
20
10
0
-10

average
SD
simulation

-20

Time (s)

-30
0

0.5

1

1.5

2

2.5

3

0

0.5

1

1.5

2

2.5

3

Figure 8-4 Hindfoot kinematics in three orientations (a, b, c) determined from motion analysis; comparisons of ankle moments
between human subject tests and model simulation in three orientations (d, e, f). Red bars indicate the SD of simulated moments.

157

From motion analysis the subjects generated approximately 27° of dorsiflexion, 12° of
eversion, and 29° of external rotation (Fig. 8-4a, 8-4b, and 8-4c, respectively) in the
hindfoot. The ankle moments from the simulations were averaged across subjects and
compared well in all three directions with those developed in experiments (Fig. 8-4d, 84e, and 8-4f).

12
§

Max Strain (%)

10

#

8
6
‡

4

‡

initial strain

2

†

*

*

0
CaFL

ATaFL

PTaFL

ATiFL

PTiFL

ADL

PDL

Ligaments
Figure 8-5 Maximum strains (averaged across subjects with ±1SD) in various ligaments
predicted with the computational model. An initial strain of 2% was applied to each
ligament. ADL: anterior deltoid ligament (ATiTL and TiNL). PDL: posterior deltoid
ligament (PTiTL and TiCL). Different symbols (§ # ‡ † *) indicate statistically
significant differences between ligaments.

158

Strains were displayed for seven ankle ligaments that have frequently been associated
with ankle injuries (Fig. 8-5). The anterior deltoid ligament (ADL) consists of two
bundles [35], the ATiTL and TiNL (Fig. 8-2). The strains generated in these bundles
were averaged. The posterior deltoid ligament (PDL) is composed of the PTiTL and
TiCL, and the strains were similarly averaged for this presentation. Strain developed in
the ATiFL (9.2 ± 1.1%) was significantly greater than all other ligaments (p<0.001).
Strain generated in the ADL (7.8 ± 0.7%) was significantly greater than the remaining
ligaments (p<0.001). Strains in the PTaFL (4.2 ± 0.5%) and the PTiFL (3.6 ± 0.7%) were
not different (p=0.093). Due to a loss of tension during external foot rotation, the CaFL,
ATaFL, and PDL experienced strains less than the initial strain of 2%. Strain in the CaFL
was different from strains in the ATaFL (p=0.013) and the PDL (p=0.011), while no
significant difference was observed between the ATaFL and the PDL (p=0.584).

159

1.2
dorsiflexion

Normalized Values

1

eversion
external rotation

0.8

ATiFL strain
ADL strain

0.6
0.4
0.2
0
Time (s)
-0.2
0

0.5

1

1.5

2

2.5

3

Figure 8-6 Temporal profiles of the averaged hindfoot motions (deg) in three orientations
(dorsiflexion, eversion and external rotation) and dynamic strains (%) in two ligaments
(ATiFL and ADL). Values were normalized between their initial and peak measures (i.e.
dorsiflexion: 9.2~26.6 deg; eversion: 1.5~12.3 deg; external rotation: 3.0~28.9 deg;
ATiFL strain: 2%~9.2%; ADL strain: 2%~7.8%).

The averaged ankle kinematics and strains in the ATiFL and ADL were normalized and
compared temporally in Fig. 8-6. Ankle dorsiflexion and eversion started prior to external
rotation during this maneuver. Strain in the ATiFL started with ankle dorsiflexion and
eversion, but it was enhanced by external rotation and peaked with dorsiflexion and
external rotation. Dorsiflexion and eversion had less influence on initiating strains in the
ADL, and strain in each decreased similarly with decreases in dorsiflexion and external
rotation.

160

8
Mean
7
dorsiflexion

6

eversion

Subjects

5

external rotation

4

ATiFL strain

3

ADL strain

2
1
0
1

1.5

2

2.5

3

3.5

Peak time (s)
Figure 8-7 Temporal sequences of peak hindfoot motions and peak ligament strains for
each subject. The mean values were shown on the top. Peak eversion occurred
statistically first, while peak strain in the ADL took place significantly later in the
external rotation maneuver.

Peak ankle eversion occurred significantly (p<0.001) prior to peak dorsiflexion, which
took place almost simultaneously with peak external rotation (Fig. 8-7). While peak strain
in the ATiFL occurred at the same time as peak dorsiflexion and external rotation
(p>0.05), peak strain in the ADL occurred significantly later (p<0.001).

161

DISCUSSION
This study estimated dynamic ankle ligament strains during external foot rotation. Results
from the OFM-based motion analysis showed dorsiflexion, eversion, and external
rotation of the hindfoot when subjects performed single-legged, internal rotation of their
bodies with a planted foot (Fig. 8-4). These motions were used to drive a generic
computational ankle model. These simulation studies supported the hypothesis that the
largest strain would be generated in the ATiFL. The ADL (ATiTL, TiNL), however, also
experienced large strains during external foot rotation simulations (Fig. 8-5). These
results may be supported by clinical [14,15,33] and biomechanical studies [13,52] that
show the deltoid ligament is often injured in severe high ankle sprains. The current study
showed ligament strains less than 10% (Fig. 8-5). These results parallel with the current
literature. A cadaver study by Ozeki et al. [39] shows that maximum strains in the ATaFL,
PTaFL, CaFL, and ATiFL for a physiological range of ankle motion are 7.9%, 5.9%,
5.3%, and 5.2%, respectively. Similarly Colville et al. [10], using cadavers with 20° of
dorsiflexion, document ankle ligament strains of approximately 4%, 6% and 8% in the
ATaFL, ATiFL and PTiFL, respectively, which have been recently simulated by Wei et
al. [50] using the current computational model. The current study also showed that peak
strains in the ATiFL occurred prior to peak strains in the ADL (Fig. 8-6). A recent
cadaver study of ankle pronation injury, showing ankle fracture and ligament rupture,
also documents medial ankle injury, largely due to deltoid ligament rupture, occurring
after ATiFL rupture [20]. These data support the observed delay of peak ADL strain in
the current study.

162

Accurate measurements of ligament strains in living humans may be crucial in
understanding ankle function, injury mechanisms, and for the optimization of treatment
and injury prevention programs. Various other studies have been conducted, using
invasive and non-invasive techniques, to measure in vivo soft tissue strains in the lower
extremity, in particular, to the anterior cruciate ligament [3,17,22,41,46], the patellar
tendon [40], the Achilles tendon [32], and the ATaFL and CaFL of the ankle [12]. While
invasive, direct measurements of ligament strain using implantable strain gauges are
valuable, they are performed under surgical procedures involving anesthesia which may
engage altered muscle function [16]. As an alternative, non-invasive methods such as the
combined dual-orthogonal fluoroscopic and MRI technique by De Asla et al. [12] have
shown promise in measuring length changes in the ATaFL and CaFL of the ankle. The
methodology, however, has only been used in static studies under non-weightbearing
conditions. The current technique, using a dynamic, rigid-body ankle model driven by
motion analysis-based kinematic data may be a good alternative to measure in vivo ankle
ligament strains dynamically and non-invasively.

Limitations of the current study should be noted. First, a generic foot model based on a
19 year old, male cadaver with a body weight of 86 kg was used in the current study.
While the six subjects recruited for the current study were from a similar demographic
population, future studies are envisioned with subject-specific models. Second, the OFM
has three foot segments (hindfoot, forefoot and hallux) and a tibial segment [9]. Since the
hindfoot segment includes the calcaneus and talus, subtalar joint motion cannot be
directly measured. This feature of the OFM may be a limitation for some ankle injury

163

simulation studies. For example, inversion injuries can create subtalar ligament lesions,
which may affect the cervical ligament, the interosseous talocalcaneal ligament, or the
lateral talocalcaneal ligament [23,24]. An additional simulation was conducted during the
current study wherein the subtalar joint was freed and four more ligaments were added to
the model: the interosseous, lateral, medial, and posterior talocalcaneal ligaments. The
results indicated negligible subtalar motion and therefore strains in these ligaments under
similar levels of external foot rotation (data not presented). Another limitation of studies
involving human subjects is the variance in subject performances. While we tried to
minimize these inconsistencies by instructing the subjects about our preferred, ideal
movement, each performed the task with some differences in magnitude of hindfoot
rotation and in a slightly different temporal response (Fig. 8-7). Such differences between
subjects may influence the results of future studies on shoe-surface interface
characteristics. Finally, while the moment-time curves from the computational model
yielded temporal characteristics similar to those of human subjects (Fig. 8-4), there was a
10-20% difference in these data. The computational model has only been validated
against cadavers [50], so the indicated difference may be due to muscles. Future studies
will be needed to include these effects in the modeling.

In conclusion, we have developed a methodology to use motion analysis data in
combination with a computational ankle model to determine ankle ligament strains
during a dynamic movement. The study showed that external foot rotation generated
large ATiFL strains, with peak strains in the ADL occurring later in the motion sequence.

164

We believe future studies could be conducted using this technique to evaluate shoesurface interface characteristics in subject-specific models.

165

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CHAPTER 9
CONCLUSION

Overview
The biomechanics of ankle injury have been studied extensively, primarily through
mechanical testing of human cadavers. Cadaveric testing is an invaluable methodology in
biomechanics, because the magnitude and direction of the loading can be measured
precisely and correlated with the resulting injury pattern. While clinical and
epidemiological studies provide useful descriptions of injury patterns that occur in the
real world, their retrospective nature precludes a definitive analysis of the forces that
caused the injury. Understanding the mechanism of ankle injuries is essential for
developing countermeasures to prevent injury and for reconstructing injurious events.
Knowledge of an injury’s mechanism can also suggest potential associated injuries,
which is helpful in diagnosis and treatment.

Computational models offer some unique insights into foot and ankle function that are
unavailable from in vivo or in vitro studies. With the appropriate computer model, any
physical variable can be measured without concern for sensors affecting the measurement.
Sensitivity studies can be performed in which one parameter of interest, and only that
parameter, is varied, eliminating potentially confounding factors such as anatomic
variations in subjects or specimens, and the passage of time. Computational simulations
are also generally faster and less expensive than experiments, which necessarily consume
physical resources. Computational modeling may be most effective, however, when

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employed in combination with experiments. The predictions of models must be tested
using real data, and models are potentially useful for both explaining experimental results
and deciding how best to allocate experimental resources. Such combinations of
modeling and experimentation have been employed in some studies described in this
dissertation, which focuses on their applications to understand the functional roles of foot
and ankle structures.

Chapter 2
Chapter 2 was a cadaveric study, wherein ten ankles were tested by externally rotating
the foot until gross injury. Foot constraint was investigated by restraining the calcaneus
tightly with potting and screws or loosely with athletic tape. Two different frequencies of
rotation (0.5 Hz and 2 Hz) were used in the study. The results showed that mean failure
torque of the ankles was 69.5 ± 11.7 Nm with a mean failure angle of 40.7 ± 7.3°. No
effects of rotation frequency or flexion angle were noted (Table 9-1). The posterior
talofibular ligament (PTaFL) was the most common injury site during external rotation of
the foot with a relatively rigid constraint. Visible damage to the syndesmosis only
occurred in combination with fibular fracture in these experiments. An anterior deltoid
ligament (ADL) injury was noted in the youngest specimen. While clinically syndesmosis
injury is often associated with external foot rotation, Chapter 2 suggested that this fact
might be due to a less constrained subtalar joint in the real-world and more typically for a
younger population. It was concluded that in future experiments the role of constraint
offered to the foot and ankle complex by footwear with varying upper rigidity should be

173

investigated as it may affect the location and torque/rotations needed to produce soft
tissue injuries in the human ankle.

Table 9-1 – Torque and foot rotation at failure from Chapter 2.

Specimen

Rotation
Freq.
(Hz)

Flexion
(deg)

Torque
(Nm)

Angle
(deg)

32388R
32388L

0.5
2.0

20
20

55
60

34
35

32416R

2.0

20

67

31

32416L

0.5

20

95

55

32516R

0.5

20

61

45

32516L

0.5

20

62

35

32489R
32489L

2.0
2.0

10
10

68
76

40
45

32532L

0.5

20

80

47

32498R

2.0

-10

71

40

AVG
(SD)

69.5
(11.7)

40.7
(7.3)

Injury Description
Fib avulsion of PTaFL
Fib avulsion of PTaFL
Distal fib fracture through
ATiFL
Distal fib fracture
Distal fib fracture through
ATiFL
Spiral fracture of tibia and
fibula
Fib avulsion of PTaFL
Fib avulsion of PTaFL
Anterior deltoid ligament
midsubstance
PTaFL midsubstance

PTaFL: posterior talofibular ligament. ATiFL: anterior tibiofibular ligament. Positive
flexion is dorsiflexion. Specimen descriptions were documented in Chapter 2.

Chapter 3
In Chapter 3 a multibody computational ankle model was developed and validated to
include the influence of foot constraint, determine the kinematics of the joint under
external foot rotation, and consequently estimate strains in various ankle ligaments. The
model included 14 rigid bones and 20 ankle ligaments formulated as linear elastic springs.
The model was used to simulate a potted foot by fusing the calcaneus to the forefoot and
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a taped foot by allowing motion of the bone. The results helped explain, by showing
ankle ligament strains (Figure 9-1), the mechanisms of injury documented in Chapter 2
that an excessive external foot rotation might generate a PTaFL injury for a rigid foot
constraint, and an ADL injury for a pliant foot constraint. It was concluded that the
computational models might be further developed and modified to simulate the human
response for different shoe designs, as well as on various athletic shoe-surface interfaces,
so as to provide a computational basis for optimizing athletic performance with minimal
injury risk.
60

Ligament Strain (%)

60

Potted Foot Model

A

Taped Foot Model

B

50

50
PTaFL
Deltoid
ATiFL

40

PTaFL
Deltoid
ATiFL

40

30

30

20

20

10

10

0

0
0

10

20

30

40

50

60

70

External Rotation (deg)

0

10

20

30

40

50

60

70

External Rotation (deg)

Figure 9-1 Strains in different ankle ligaments during external rotation of the
computational foot models: (A) potted foot; (B) taped foot. For clarity, only strains of the
ligaments injured in the cadaver experiments (Chapter 2) were plotted. Detailed results
from these figures were documented in Chapter 3.

175

Chapter 4
Chapter 4 includes two simulation studies, using the generic ankle model developed in
Chapter 3, on the mechanisms of injury in a lateral ankle sprain and a high ankle sprain.
In order to simulate an injury event that had accidentally occurred in a laboratory and was
documented as a grade I anterior talofibular ligament (ATaFL) sprain, approximately
three times body load was applied to the ankle model, and three motion scenarios were
simulated: (1) a pure inversion of 48°; (2) a 48° inversion plus a 20° plantarflexion; and
(3) a combination of 48° inversion, 20° plantarflexion, and 30° internal rotation.
Ligament strains were compared for the three motion scenarios (Figure 9-2). For pure
inversion, the CaFL strained the most at 12%, and the ATaFL and the LTaCL were in
lower at about 10%. For the scenario of inversion, plantarflexion, and internal rotation,
the highest strain occurred in the ATaFL at 20.5%, followed by the CaFL at 16%. The
study concluded that, in addition to inversion, ankle plantarflexion and internal rotation
play important roles in an ankle inversion injury. The highest strain seen in the ATaFL
with a combination of inversion, plantarflexion, and internal rotation was in concert with
the motions documented in the accidental ankle injury.

176

25
Inv.

Strain (%)

20

Inv.+ Plan.Flex.
Inv.+ Plan.Flex.+ Int.Rot.

15
10
5

aC
L
PT

L
IT
aC

C
L
LT
a

aF
L
PT

Ta
FL
A

aF
L
C

iF
L
PT

A

Ti
FL

0

Ligaments

Figure 9-2 Strains in various ligaments for different ankle motions. Eight ligaments were
selected based on strain higher than 2% for at least one scenario. Those ligaments are the
anterior and posterior tibiofibular (ATiFL and PTiFL); the calcaneofibular (CaFL); the
anterior and posterior talofibular (ATaFL and PTaFL); and the lateral, interosseous, and
posterior talocalcaneal (LTaCL, ITaCL, and PTaCL).

In order to simulate several injury cases reported and diagnosed as high ankle sprains in
NCAA football games, approximately three times body weight was applied to the ankle
model, and four motion scenarios were simulated: (1) a pure external rotation of 30°; (2)
a 30° external rotation plus a 20° dorsiflexion; (3) a 30° external rotation plus a 20°
eversion; and (4) a combination of 30° external rotation, 20° dorsiflexion, and 20°
eversion. Ligament strains were compared (Figure 9-3). Under a pure external rotation,
the ATiTL strained the most at 8%, while the PTaFL and the ATiFL were strained about
6% and 3.5%, respectively. Dorsiflexion and eversion, added to the external rotation,
significantly increased ATiFL strain to 9% and 17%, respectively. For the combination of
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all three motions, the highest strain occurred in the ATiFL at 21%, followed by the
ATiTL 15.5%. The study concluded that external rotation, dorsiflexion and eversion each
played an important role in a high ankle sprain. The highest strain observed in the ATiFL
occurred with a combination of all three motions, in concert with the motions seen in
game films.

These simulation studies may add to the knowledge about the potential mechanisms of
ankle sprains and provide a computational basis for studies of other ankle injuries or
ankle sprain prevention strategies.
25

Ext.Rot.
Ext.Rot.+ Dor.Flex.
Ext.Rot.+ Eve.
Ext.Rot.+ Dor.Flex.+ Eve.

Strain (%)

20
15
10
5
0
Inte-II

ATiFL

PTiFL

ATiTL

TiNL

PTaFL

Ligaments

Figure 9-3 Strains in various ligaments for different ankle motions. Six ligaments were
selected based on strain being higher than 5% for at least one motion. Inte-II is the
interosseous ligament II, ATiTL is the anterior tibiotalar ligament, and TiNL is the
tibionavicular ligament.

178

Chapter 5
Chapter 5 was a cadaver study with an objective to produce, for the first time, a high
ankle sprain in a laboratory. The study restrained pairs of cadaver ankles using elastic
athletic tape, externally rotated axially pre-loaded limbs in dorsiflexion with and without
foot eversion, and identified isolated ligamentous injury to the ATiFL in the everted
limbs as opposed to deltoid ligament injury in the neutrally positioned limbs (Table 9-2).
In addition, rotation at failure for everted limbs was significantly greater than for neutral
limbs (46.8 ± 6.1° versus 37.7 ±5.4°). The failure torques, however, were not statistically
different. The study concluded that external rotation of a highly everted foot would
generate high ankle sprains. Eversion of an axially loaded foot predisposed the ATiFL to
injury, which may describe a clinical basis for high ankle sprains. Despite its
biomechanical nature, the study may heighten clinical awareness of isolated ATiFL
injury in cases of foot eversion and dorsiflexion prior to external rotation, as opposed to
deltoid ligament injury in cases of foot dorsiflexion followed by external rotation, and
provide guidance to help study the effect of prophylactic measures, such as special shoe
designs or ankle braces to limit excessive foot eversion.

179

Table 9-2 Foot rotation at failure, and resultant injury.

Specimen

Failure rotation
(deg)
Neutral Everted

1

36.0

43.4

2

47.1

55.3

3

35.1

50.1

4

39.8

50.0

5

31.1

43.4

6

37.0

38.6

Mean

37.7

46.8*

SD

5.4

6.1

Note:

Failure mode
Neutral
Tibial avulsion of
the ADL
Partial tear of the
ADL
Partial tear of the
ADL and ATiFL
Rupture of the
deltoid ligament
Spiral fracture of
the fibula
Rupture of the
ADL

Everted
Rupture of the ATiFL
Partial tear of the ATiFL
Rupture of the ATiFL
Rupture of the ATiFL
Partial tear of the ATiFL
Rupture of the ATiFL

The posterior tibiofibular ligament and the
interosseous ligament remained intact in all
specimens.

The right limbs were neutral in the coronal plane, while the left limbs were
everted 20°.
* Statistically different than the neutral foot.

Chapter 6
Chapter 6 was a cadaver study that investigated the effects of rotational stiffness of
football shoes on talus motion and ankle ligament strain patterns during external rotation
of the foot. Four football shoe designs were tested and compared in terms of rotational
stiffness. Twelve (6 pairs) male, cadaver lower extremity limbs were externally rotated
30º using two selected football shoe designs, i.e. a flexible shoe and a rigid shoe. Motion
capture was used to track the movement of the talus with a reflective marker array
screwed into the bone. A computational ankle model was utilized to input talus motions
for the estimation of ankle ligament strains. At 30º of rotation the rigid shoe generated
180

higher ankle joint torque at 46.2 ± 9.3 Nm than the flexible shoe at 35.4 ± 5.7 Nm. While
talus rotation was greater in the rigid shoe (15.9 ± 1.6º vs. 12.1 ± 1.0º), the flexible shoe
generated more talus eversion (5.6 ± 1.5º vs. 1.2 ± 0.8º). While these talus motions
resulted in the same level of ADL strain (approximately 5%) between shoes, there was a
significant increase of ATiFL strain (4.5 ± 0.4% vs. 2.3 ± 0.3%) for the flexible versus
more rigid shoe design (Figure 9-4). The flexible shoe may provide less restraint to the
subtalar and transverse tarsal joints, resulting in more eversion but less axial rotation of
the talus during foot/shoe rotation. The increase of strain in the ATiFL may have been
largely due to the increased level of talus eversion documented for the flexible shoe. The
study may provide some insight into relationships between shoe design and ankle
ligament strain patterns. In future studies these data may be useful in characterizing shoe
design parameters and balancing potential ankle injury risks with player performance.

Figure 9-4 The horizontal bar indicates significant difference between shoes. The strain
in the ADL was statistically greater than in the ATiFL for both shoe designs.

181

Chapter 7
Chapter 7 modified the previously developed, generic ankle model to investigate shoe
and shoe-surface interface effects on ankle ligament strains during a simulated sidestep
cutting task. Briefly, a compressive pre-load of two times body weight was distributed
between the tibia and the fibula. Two kinds of foot constraint were simulated,
representing rigid and pliant shoes. The rigid shoe included a fused subtalar joint, while
the pliant shoe allowed the motion between the talus and the calcaneus. The shoe-surface
interface was simulated with a torsional spring at the center of a surface that was inserted
under the foot, to represent the mechanical effect of a playing surface. Two kinds of
surfaces, synthetic turf and natural grass, were simulated with stiffnesses of 3.6 and 2.2
Nm/deg, respectively, based on the literature. A rotary motor was applied to the surface
and aligned at the ankle center between the malleoli. The external rotation angle was set
to 30° to simulate a common sidestep cutting task. The results showed that the maximum
strains in key ligaments were larger on the synthetic turf than on natural grass. The order,
from the highest to the lowest strain, was rigid shoe on turf, pliant shoe on turf, rigid shoe
on grass, and pliant shoe on grass (Figure 9-5). The study indicated that shoe-surface
interface influenced ankle ligament strains and consequently the potential for ankle injury.
The maximum strain occurred in the PTaFL for the rigid shoe condition, while the deltoid
ligament experienced the highest strain in the pliant shoe, suggesting that the location of
ankle injury might also depend on foot constraint. Ultimately, such computational models
may provide a basis for optimizing shoe designs and shoe-surface interface
characteristics to enhance player performance and minimize injury risk.

182

Ligament Strain (%)

40

Rigid / Turf

40

Pliant / Turf

40

Rigid / Grass

40

Pliant / Grass

30

30

30

30

20

20

20

20

10

10

10

10

0
40 0

0
40 0

0
40 0

0
0

10

20

30

External Rotation
(deg)

10

20

30

External Rotation
(deg)

10

20

30

External Rotation
(deg)

PTaFL
Deltoid
ATiFL

10

20

30

40

External Rotation
(deg)

Figure 9-5: Strains in various ligaments during external rotation of the surface for
different shoes and shoe-surface interface conditions.

Chapter 8
Chapter 8 was an in vivo study to determine dynamic ankle ligament strains from a
computational model driven by motion analysis based kinematic data. Briefly, six
subjects performed single-legged, internal rotation of the body with a planted foot while a
marker-based motion analysis was conducted to track the hindfoot motion relative to the
tibia. These kinematic data were used to drive an established computational ankle model.
Ankle ligament strains, as a function of time, were determined. The ATiFL experienced
the highest strain at 9.2 ± 1.1%, followed by the ADL at 7.8 ± 0.7%, averaged over the
six subjects (Figure 9-6). The peak ATiFL strain occurred prior to peak strain in the ADL
in all cases. This novel methodology may provide new insights into mechanisms of high
ankle sprains and offer a basis for future evaluations of shoe-surface interface
characteristics using human subjects rather than mechanical surrogate devices.

183

12
§

Max Strain (%)

10

#

8
6
‡

4

‡

initial strain

2

†

*

*

0
CaFL

ATaFL

PTaFL

ATiFL

PTiFL

ADL

PDL

Ligaments
Figure 9-6 Maximum strains (averaged across subjects with ±1SD) in various ligaments
predicted with the computational model. An initial strain of 2% was applied to each
ligament. PDL is the posterior deltoid ligament. Different symbols (§ # ‡ † *) indicate
statistically significant differences between ligaments.

184

Summary
In summary, this dissertation has offered three major contributions to the current
literature.

(1) High ankle sprain mechanisms
The mechanisms of injury for high ankle sprains can be confusing because of the
different anatomic structures involved and the manner in which these structures develop
excessive stress in the three planes of motion. A primary contribution of this dissertation
to the current literature was the clarification of a high ankle sprain mechanism. Clinically,
the mechanism of high ankle sprain has mostly been attributed to external foot rotation.
But no experimental study documents this injury under such a loading condition. In
Chapter 5 we have produced, for the first time, an isolated ligamentous injury to the
anterior tibiofibular ligament with the human cadaver foot by external rotation of a highly
everted, dorsiflexed foot. In contrast, external rotation of a dorsiflexed foot but without
eversion generates a deltoid ligament injury. The study, from a biomechanical point of
view, clarifies that eversion of an axially loaded foot predisposes the anterior tibiofibular
ligament to injury, forming a basis for a high ankle sprain. The study also shows that foot
eversion is the key in transferring injury from the deltoid ligament to the anterior
tibiofibular ligament. A better understanding of a mechanism of injury may ensure
developments of new prevention strategies, preferable treatment guidelines, and
improved rehabilitation programs for the injury.

185

Limitations of the biomechanical studies include unrealistic foot restraint, lack of muscle
effects, limited testing speed, potential micro-damage to soft tissues due to repeated
testing, etc. Future physiological studies involving human subjects could be conducted as
a continuation to investigate mechanism, prevention, and/or rehabilitation of high ankle
sprains with in vivo settings.

(2) Computational modeling of the ankle joint
A second major contribution of this dissertation to the current literature has been the
development and validation of a generic computational ankle model that was constructed
from a computed tomography (CT) scan of a cadaver foot/ankle complex (Chapter 3). CT
images were first converted into 3D models in MIMICS and then imported into rigidbody motion simulation software SolidWorks. The ankle model includes 21 ligaments
formulated as linear elastic springs with properties adapted from the literature. In addition
to Chapter 3, wherein the model was applied to simulate potted and taped foot restraints,
it was utilized in Chapter 4 to simulate lateral and high ankle sprain cases and in Chapter
7 to simulate a sidestep cutting task. This model was also used in combination with in
vitro (Chapter 6) and in vivo (Chapter 8) experiments to help estimate ankle ligament
strains by input of those collected, bone motion data.

Limitations of the computational ankle model include its generic nature (non-subjectspecific), ligaments being simulated as linear springs, and lack of muscle and cartilage
elements. Subject-specific models could be developed from CT or MRI images of
cadavers or human subjects in future studies. In fact, the cadaver limbs used in Chapter 5

186

have been scanned and the images are being used to build subject-specific models for
simulation purposes in an ongoing study. Results from that study are not included in this
dissertation.

(3) Investigation of shoe properties
Shoe-surface interface characteristics have been implicated in the high incidence of ankle
injuries. However, the differences in rotational stiffness between shoes may also
influence injury risk. Another major contribution of this dissertation to the current
literature is the initiation of study on the influence of various shoe designs in producing
ankle ligament strains. In Chapter 6 we documented the rotational stiffness of four
football shoe designs, and then externally rotated six pairs of cadaver limbs in two
different shoes. The study showed that football shoe design can have an effect on the
pattern of ankle ligament strains during external rotation of the foot to potentially
influence the location and severity of a subsequent ankle injury.

While that study was indeed limited to a subfailure testing level, it represents a first step
in attempt to understand the effect of football shoe design on the potential for ankle
injury. Future studies could involve failure tests or investigate shoe-surface interface
characteristics using the same experimental model.

187

APPENDIX

188

APPENDIX

Standard Operating Procedures (SOP) for Computational Modeling of the Ankle
This SOP describes how to develop a computational foot model from a CT scan
using MIMICS and SolidWorks.
Before you start, make sure the foot being CT scanned in an expected position:
neutral or certain degrees of dorsiflexion/plantarflexion etc.
This SOP is based on a generic CT scan of the foot. If needed, one can also build
specimen-specific foot models using similar procedures.
This modeling approach uses MIMICS 12.1v (Materialise, Ann Arbor, MI) and
SolidWorks 2009 x64 Edition (TriMech Solutions, LLC, Columbia, MD).
Several references are extremely useful in helping the model development. They
are listed at the end of this document.

Preparation of CT scans:
1. A fresh frozen, male, left cadaver foot (19 years old), showing no signs of
abnormal anatomy, was transected approximately 15 cm distal to the knee center.
2. The foot was fully thawed overnight and CT scanned to obtain three-dimensional
joint anatomy.
3. The CT scan, spanning the tibia to foot, was performed in MSU Radiology
Department.
4. While scanning, the foot was in a neutral anatomical orientation (approximately
90 degrees between the foot plane and the lower leg).

189

5. The CT scan provided the individual slices from which the three-dimensional
anatomy was reconstructed. CT images for the shank and foot were available in
0.6 mm intervals.

CT scan Æ 3D model by MIMICS:
1. Detailed features of the foot/ankle complex were obtained by importing Digital
Imaging and Communications in Medicine (DICOM) files of individual CT
images into Materialise’s Interactive Medical Imaging Control System (MIMICS).
2. In MIMICS, a mask for each bone of the foot/ankle complex was created. To
create each mask, thresholding was performed to achieve a maximum contrast
between the bones (white) and the surrounding soft tissues (gray and black).
3. The area representing each bone of interest was then selected and covered with a
specified color mask on the entire stack of CT images (Figure A-1).
4. Fourth order polynomials approximated the bone perimeter on each scan. This set
of perimeters was used to create three-dimensional surface models for each bone.
5. For some bones that possessed multiple perimeters in several consecutive CT
images, such as the head, lateral process, and posterior process of the distal talus,
two surface files had to be created.
6. The surfaces for each bone were then saved as Stereolithography (STL) files for
export.
7. To reduce the size of the surface files and the subsequent models, STL files were
remeshed in MIMICS to smooth the surface of each bone.

190

8. The size of the surface files after remesh should be less than 100 KB for the best
post-process operations.
9. The remeshed STL files were then imported into the 3D solid modeling program
SolidWorks to render as solid objects.
10. An example of MIMICS interface was shown below for reference.

Figure A-1: A typical MIMICS interface. Each bone was assigned to a different color.

Three-dimensional reconstruction of the foot in SolidWorks:
1. In SolidWorks, the ScanTo3D package was used to reconstruct each bone and
simplify the bone surfaces.
2. The separate solid objects within SolidWorks representing the bones were
assembled together to form the foot/ankle complex.
3. Due to the curve and surface fitting techniques, a small amount of overlap could
be present between bones (very unlikely).

191

4. Overlapping could be resolved using the cavity feature within SolidWorks to
create a cavity in the bone with the concave geometry. The cavities were made
slightly larger than the intersecting portion of the convex bone (scaled by 0.01%)
to create a small space between the two bones.
5. To remedy the sharp edges of each cavity, a fillet could be created around each
cavity’s perimeter. No fillet’s radius exceeded 0.5 mm.
6. The three-dimensional SolidWorks model of the lower extremity was then
transferred into SolidWorks Motion, a software package designed for mechanical
system simulation.
7. SolidWorks Motion (SolidWorks, TriMech Solutions, LLC, Columbia, MD) was
subsequently implemented to apply ligamentous restraints, prescribe force/motion
constraints, and simulate model dynamics.
8. Soft tissue reconstruction of ligamentous geometry was achieved through
SolidWorks Motion simulation elements. These elements were used to describe,
individually and in arrays, the major dorsal, plantar, medial, lateral, and
interosseous ligaments that constrain the leg, ankle, hindfoot, midfoot, and
forefoot bones, maintaining the structure of the foot/ankle complex and leg.
9. Multi-element arrays could be used to describe ligaments with broad insertion
sites and complex geometry in multiple planes in order to capture the threedimensional rotational and translational control these ligaments impart on their
respective bones. The geometry of these arrays and all ligaments attachment sites
were taken from reported dissection of foot and ankle ligaments in literature,

192

supplemented with anatomy text and in house dissection for additional
clarification.
10. Ligaments were represented as linear spring elements, with stiffness values taken
from the literature. Ligaments preloads were induced by reducing lengths by 2%.
11. Each bone was allowed to move in all six degrees of freedom, leaving body
motion to be a function of ligament behavior, muscle forces, surface contact, and
external perturbations.
12. The 3D articular contact conditions were simulated between adjacent bone
surfaces as well as contact between the bones and ground plate (if there is any).
For contact behavior, COSMOSMotion applied interference detection to
determine regions of overlap between adjacent surfaces. Using the centroid of
each interfering volume as the point of contact, equal and opposite reaction forces
were then applied to each body along the contact normal direction.
13. Friction and the effects of gravity were considered negligible.
14. Simulated body load was applied to the proximal end of the model, proportionally
distributing it between the tibia and the fibula as to one-sixth loading on the fibula.

References:
1. Feng Wei, Stanley C. Hunley, John W. Powell, Roger C. Haut. Development and
validation of a computational model to study the effect of foot constraint on ankle
injury due to external rotation. Annals of Biomedical Engineering 39(2):756-765,
2011.

193

2. Peter C. Liacouras, Jennifer S. Wayne. Computational modeling to predict
mechanical function of joints: application to the lower leg with simulation of two
cadaver studies. Journal of Biomechanical Engineering 129:811-817, 2007.
3. Joseph M. Iaquinto, Jennifer S. Wayne. Computational model of the lower leg and
foot/ankle complex: application to arch stability. Journal of Biomechanical
Engineering 132:021009, 2010.
4. Justin P. Fisk, Jennifer S. Wayne. Development and validation of a computational
musculoskeletal model of the elbow and forearm. Annals of Biomedical
Engineering 37(4):803-812, 2009.
5. Edward M. Spratley, Jennifer S. Wayne. Computational model of the human
elbow and forearm: application to complex varus instability. Annals of
Biomedical Engineering 39(3):1084-1091, 2011.

194

Foot Rotations Calculated in a Joint Coordinate System
The following schematics (Figure A-2) and equations show calculations for the clinical
description of foot rotations in a Joint Coordinate System (JCS), namely, foot inversion /
eversion, internal / external rotation, and dorsiflexion / plantar flexion. Briefly, if the
initial and final positions of the foot are described in X-Y-Z and X'-Y'-Z' coordinates,
respectively, the JCS is then established by a non-orthogonal coordinate e1-e2-e3, where
e2 is parallel with Y, e3 is parallel with Z', and e1 is the cross product of e2 and e3. Foot
rotations are calculated by the dot products of these vectors (shown in the equations). A
couple of references are listed after Figure A-2.

Figure A-2: Schematics showing the calculations of foot rotation angles in a Joint
Coordinate System.

195

References:
1. Grood ES, Suntay WJ. A joint coordinate system for the clinical description of
three-dimensional motions: application to the knee.

J Biomech Eng.

1983;105(2):136-144.
2. Soutas-Little RW, Beavis GC, Verstraete MC, Markus TL. Analysis of foot
motion during running using a joint co-ordinate system. Med Sci Sports Exerc.
1987;19(3):285-293.

196

Table A-1: Raw Data from the Cadaver Failure Tests (Chapter 5)

03952LFailure50deg

33438Lfailure60deg

33449Lfailure80deg

Failure

Failure

Failure

Torque Nm Rotation deg Torque Nm Rotation deg Torque Nm Rotation deg
78.30

43.38

91.83

increment

55.34

94.51

increment

50.08

increment

0.00

0.00

0.00

0.00

0.00

0.00

-1.62

5.00

3.08

4.99

-0.91

5.06

2.47

9.99

7.95

10.03

0.67

10.14

8.39

15.01

12.35

14.91

1.90

15.15

16.83

20.07

18.47

19.92

3.22

20.08

27.47

24.98

27.21

25.07

18.05

25.01

41.15

30.10

37.20

30.00

37.06

30.24

56.83

35.09

48.86

35.14

58.79

35.18

71.96

39.98

60.46

39.92

80.60

40.10

78.30

43.38

73.40

45.10

82.98

45.09

46.18

44.98

85.09

50.06

94.51

50.08

54.64

50.00

91.83

55.34

86.41

59.99

197

Table A-1 (cont’d)
33477Lfailure40deg

33608LFailure50deg

33686LFailure40deg

Failure

Failure

Failure

Torque Nm Rotation deg Torque Nm Rotation deg Torque Nm Rotation deg
103.75

50.04

87.18

increment

43.40

86.51

increment

38.64

increment

0.00

0.00

0.00

0.00

0.00

0.00

5.10

5.05

-0.14

5.02

1.60

5.04

11.57

10.01

5.32

10.03

9.46

10.08

18.16

15.06

12.33

15.12

17.44

15.03

24.36

20.10

20.57

20.09

28.81

20.01

32.97

25.02

30.72

25.01

42.59

25.02

43.91

30.15

44.27

30.10

59.12

30.10

57.36

35.10

60.25

35.05

75.60

35.00

73.18

40.12

77.16

40.03

86.51

38.64

90.31

45.12

87.18

43.40

75.51

40.00

103.75

50.04

81.70

45.11

52.06

55.10

83.36

50.00

53.10

60.03

198

Table A-1 (cont’d)
03952RFailure40deg

33438Rfailure50deg

33449Rfailure80deg2

Failure

Failure

Failure

Torque Nm Rotation deg Torque Nm Rotation deg Torque Nm Rotation deg
87.42

36.02

80.87

increment

47.06

87.98

increment

35.09

increment

0.00

0.00

0.00

0.00

0.00

0.00

1.87

5.03

1.06

5.02

5.92

5.02

8.94

9.97

5.80

10.01

17.15

10.09

19.54

14.98

11.50

15.07

32.87

15.11

33.13

20.09

18.37

20.01

52.77

20.01

50.96

25.09

27.43

25.08

74.14

24.98

71.14

29.96

38.28

30.01

73.69

30.19

85.79

35.09

51.32

35.03

87.98

35.09

87.42

36.02

65.60

40.10

65.90

40.03

76.72

45.03

80.87

47.06

71.55

50.01

199

Table A-1 (cont’d)
33477Rfailure60deg

33608RFailure40deg

33686RFailure40deg

Failure

Failure

Failure

Torque Nm Rotation deg Torque Nm Rotation deg Torque Nm Rotation deg
101.46

39.83

53.13

increment

31.14

95.29

increment

37.00

increment

0.00

0.00

0.00

0.00

0.00

0.00

7.58

5.05

3.87

5.06

4.33

5.02

18.91

10.12

11.36

10.09

15.05

10.02

30.99

15.08

19.50

15.04

27.88

15.03

45.12

20.06

28.78

20.04

42.67

20.08

59.92

25.12

38.61

25.02

59.06

25.00

75.16

30.16

50.48

30.00

75.25

30.10

89.83

35.03

53.13

31.14

89.98

35.05

101.46

39.83

47.13

35.08

95.29

37.00

47.78

40.01

85.21

40.08

200

Table A-2: Raw Data from the Cadaver Subfailure Tests with Shoes (Chapter 6)

03952LAir30deg

33438Lair30deg

33449Rsuperbad30deg2

Torque Nm Rotation deg Torque Nm Rotation deg Torque Nm Rotation deg
0.00

0.00

0.00

0.00

0.00

0.00

-1.54

5.00

1.24

5.00

0.65

5.00

3.43

10.00

5.89

10.00

-2.21

10.00

10.91

15.00

12.16

15.00

-4.78

15.00

20.16

20.00

19.96

20.00

-8.43

20.00

29.76

25.00

28.97

25.00

-12.94

25.00

38.51

30.00

36.89

30.00

-15.81

30.00

33477LAir30deg2

33608Lair30deg

33686LAir30deg

Torque Nm Rotation deg Torque Nm Rotation deg Torque Nm Rotation deg
0.00

0.00

0.00

0.00

0.00

0.00

-6.66

5.00

-4.22

5.00

-4.54

5.00

-0.30

10.00

2.11

10.00

2.38

10.00

7.80

15.00

7.26

15.00

10.15

15.00

17.30

20.00

13.27

20.00

19.20

20.00

27.05

25.00

19.55

25.00

28.80

25.00

37.87

30.00

25.25

30.00

38.43

30.00

201

Table A-2 (cont’d)
03952RFlyposite30deg

33438RFlyposite30deg2

33449LFlyposite30deg

Torque Nm Rotation deg Torque Nm Rotation deg Torque Nm Rotation deg
0.00

0.00

0.00

0.00

0.00

0.00

0.12

5.00

3.19

5.00

-5.17

5.00

8.70

10.00

8.76

10.00

-2.18

10.00

20.29

15.00

15.88

15.00

0.61

15.00

34.16

20.00

24.68

20.00

4.20

20.00

48.56

25.00

34.08

25.00

7.52

25.00

58.41

30.00

42.41

30.00

11.63

30.00

33477RFlyposite30deg

33608Rflyposite30deg

33686RFlyposite30deg

Torque Nm Rotation deg Torque Nm Rotation deg Torque Nm Rotation deg
0.00

0.00

0.00

0.00

0.00

0.00

-3.67

5.00

1.51

5.00

3.22

5.00

2.47

10.00

8.05

10.00

10.95

10.00

10.85

15.00

15.15

15.00

20.11

15.00

21.11

20.00

22.58

20.00

31.34

20.00

32.40

25.00

29.52

25.00

42.43

25.00

42.85

30.00

34.80

30.00

52.68

30.00

202

Table A-3: SolidWorks Input and Output (Chapter 8)

INPUT – Rotation Angle (deg)
Dorsiflexion (+)

Inversion (+)

Internal rotation (+)

Plantarflexion (-)

Eversion (-)

External rotation (-)

0.00

9.27

-1.60

-3.00

0.04

9.29

-1.60

-2.37

0.08

9.31

-1.62

-2.03

0.12

9.32

-1.65

-2.02

0.16

9.33

-1.69

-2.21

0.20

9.32

-1.73

-2.47

0.24

9.30

-1.78

-2.68

0.28

9.29

-1.83

-2.80

0.32

9.29

-1.89

-2.87

0.36

9.30

-1.94

-2.87

0.40

9.33

-1.97

-2.81

0.44

9.37

-1.97

-2.66

0.48

9.42

-1.94

-2.45

0.52

9.47

-1.87

-2.21

0.56

9.52

-1.76

-1.95

0.60

9.56

-1.61

-1.66

0.64

9.58

-1.44

-1.34

0.68

9.55

-1.26

-1.03

Time (s)

203

Table A-3 (cont’d)
0.72

9.50

-1.10

-0.81

0.76

9.43

-1.00

-0.79

0.80

9.38

-0.96

-1.00

0.84

9.39

-1.00

-1.43

0.88

9.51

-1.14

-1.99

0.92

9.76

-1.35

-2.53

0.96

10.11

-1.62

-2.95

1.00

10.51

-1.91

-3.21

1.04

10.91

-2.22

-3.33

1.08

11.28

-2.55

-3.38

1.12

11.65

-2.95

-3.42

1.16

12.04

-3.48

-3.53

1.20

12.53

-4.18

-3.77

1.24

13.15

-5.05

-4.17

1.28

13.91

-6.07

-4.71

1.32

14.79

-7.18

-5.39

1.36

15.78

-8.30

-6.24

1.40

16.86

-9.38

-7.32

1.44

18.03

-10.35

-8.76

1.48

19.31

-11.17

-10.58

1.52

20.69

-11.79

-12.68

1.56

22.04

-12.09

-14.91

204

Table A-3 (cont’d)
1.60

23.23

-11.97

-17.20

1.64

24.18

-11.47

-19.65

1.68

24.96

-10.78

-22.28

1.72

25.65

-10.10

-24.83

1.76

26.25

-9.52

-26.91

1.80

26.69

-8.99

-28.23

1.84

26.81

-8.40

-28.67

1.88

26.51

-7.60

-28.27

1.92

25.76

-6.56

-27.16

1.96

24.63

-5.35

-25.48

2.00

23.33

-4.18

-23.42

2.04

22.03

-3.28

-21.18

2.08

20.84

-2.81

-18.96

2.12

19.79

-2.84

-16.96

2.16

18.83

-3.32

-15.35

2.20

17.92

-4.11

-14.21

2.24

16.99

-5.00

-13.44

2.28

15.96

-5.74

-12.79

2.32

14.84

-6.19

-12.03

2.36

13.71

-6.29

-11.13

2.40

12.72

-6.11

-10.23

2.44

11.98

-5.67

-9.49

205

Table A-3 (cont’d)
2.48

11.49

-5.01

-8.96

2.52

11.21

-4.12

-8.51

2.56

11.04

-3.02

-7.91

2.60

10.94

-1.77

-6.98

2.64

10.89

-0.51

-5.73

2.68

10.88

0.63

-4.31

2.72

10.92

1.52

-2.94

2.76

10.99

2.07

-1.78

2.80

11.06

2.27

-0.87

2.84

11.12

2.15

-0.20

2.88

11.16

1.78

0.28

2.92

11.18

1.26

0.56

2.96

11.16

0.70

0.63

3.00

11.11

0.17

0.50

OUTPUT – Resistive Moment (Nm)
Dorsiflexion (+)

Inversion (+)

Internal rotation (+)

Plantarflexion (-)

Eversion (-)

External rotation (-)

0.00

7.54

0.04

5.37

0.10

6.92

0.17

5.37

0.20

7.32

-0.20

5.60

0.30

7.25

-0.10

5.65

Time (s)

206

Table A-3 (cont’d)
0.40

6.79

0.14

5.71

0.50

6.46

0.10

5.60

0.60

6.66

-0.09

5.50

0.70

7.19

-0.05

5.24

0.80

8.25

-0.30

5.50

0.90

8.25

-0.38

5.50

1.00

8.53

-0.93

5.50

1.10

9.53

-1.76

6.00

1.20

10.29

-3.08

8.34

1.30

10.17

-4.68

10.09

1.40

8.41

-5.62

7.77

1.50

7.25

-6.00

2.25

1.60

6.53

-5.53

-8.38

1.70

6.05

-4.46

-14.83

1.80

8.97

-3.00

-17.31

1.90

18.30

-1.62

-14.11

2.00

22.83

-1.21

-9.57

2.10

19.74

-1.73

-3.16

2.20

12.45

-2.36

1.27

2.30

11.21

-2.03

2.66

2.40

8.78

-0.93

5.03

2.50

7.36

0.78

7.36

207

Table A-3 (cont’d)
2.60

3.72

2.35

1.22

2.70

1.84

3.18

1.06

2.80

1.72

2.57

1.00

2.90

1.84

1.25

1.37

3.00

2.28

0.92

0.96

208