A study of plantar stresses underneath metatarsal heads in the human foot 1

A STUDY OF PLANTAR STRESSES UNDERNEATH
METATARSAL HEADS IN THE HUMAN FOOT






CHEN, WEN-MING









NATIONAL UNIVERSITY OF SINGAPORE
2011
A STUDY OF PLANTAR STRESSES UNDERNEATH
METATARSAL HEADS IN THE HUMAN FOOT






CHEN, WEN-MING
(B.Sc (Double Degree), M.Eng)

A THESIS SUBMITTED
FOR THE DEGREE OF DOCTOR OF PHILOSOPHY


DIVISION OF BIOENGINEERING
NATIONAL UNIVERSITY OF SINGAPORE
2011

1
ACKNOWLEDGMENTS
This work was supported by a grant entitled, ‘TDSI/09-009/1A: development of
next generation combat boots with reduced shear force mechanism’ from the
Temasek Defense Systems Institute, and I am also very grateful to NUS Faculty
of Engineering Scholarship which financially supported me during my four-year
doctoral studies in Singapore.
The work would not have been possible without continuous guidance and
support from my thesis committee members. I would like to express here my
sincere respect and deepest gratitude to Dr. Taeyong Lee, Prof. James Goh, and
Assoc/Prof. Toh SL. I would particularly like to thank Prof. Victor Shim from
Department of Mechanical Engineering NUS, for the use of his lab facilities on
strain-gauging, DAQ (oscilloscope), as well as the numerous helpful discussions
during our monthly meetings. Special thanks are due also to my previous mentor
Prof. Sung-Jae Lee, from INJE University, and Assoc/Prof Peter Lee, from
University of Melbourne, for taking intense academic interest in this study as well
as encouraging me throughout these years. Sincere thanks are extended to our
clinical collaborator, Dr. Adriaan Erasmus from Rehabilitation Center, National
University Hospital, for many valuable comments which indeed helped improve
this thesis.

Finally, I would like to express my loving thanks to my wife Miao who
accompanied me at NUS through every hardship that I have encountered in the
past few years. Without her understanding, encouragement and support it would
not have been possible for me to finish this work. I love you, and thank you.

2
TABLE OF CONTENTS
List of Tables………………………………………………………………….…………7
List of Figures…………………………………………………………… ….………….9
List of Abbreviations and Symbols………………………………………………… 14
Chapter 1 Introduction
1.1. Biomechanics of neuropathic diabetic foot ulcers………………………… 17
1.2. Literature review……………………………………………………………… 19
1.2.1. Plantar stresses and mechanical etiology of diabetic foot ulcers…19
1.2.2. Challenges in measuring plantar shear stresses……………………21
1.2.3. Modeling and simulation of foot mechanism for three-dimensional
interfacial stress predictions…………………………………………….… 24
1.3. Objectives and overview of thesis……………………………………………27

Chapter 2 Numerical modeling of human foot
2.1 Finite element model of the foot-ankle complex……………………………30
2.1.1 Finite element method…………………………………………………30
2.1.2 3-D reconstruction of foot geometry…………………………………31
2.1.3 Model discretization (element selection) ……………………………33
2.2 Material properties for finite element modeling………………………………34
2.2.1 Cortical and cancellous bones, ligaments, and cartilages………….34
2.2.2 Achilles and other flexor tendons…………………………………… 35
2.2.3 Constitutive model for plantar soft tissue………………………… …38
2.2.3.1 Hyperelastic material models………………… ……………38
2.2.3.2 Ogden model………………………………………………….40

2.2.4 Material model for foot-support surface………………………………41
2.2.4.1 Insoles as supporting interface for foot…………………… 41
2.2.4.2 ABAQUS
®
hyperfoam material model………………………43
2.3 Loading and boundary conditions at push-off………………………………44
2.3.1 Articular joint movement ………………………………………………45
2.3.2 Plantar flexor muscle forces ………………………………………….47

3
2.3.2.1 Roles of the extrinsic plantar flexor muscles……………….47
2.3.2.2 Calculation of musculoskeletal loads……………………… 48
2.3.3 Foot-ground interaction with frictional contact……………………… 51
2.3.3.1 Coulomb friction…………………………………………… 51
2.3.3.2 Contact problems at foot support interface………… ……53
2.4 Finite element foot model outputs……………………………………………54

Chapter 3 Experimental technique
3.1 Instrumented soft tissue tester…………………………………………………57
3.1.1 Measurement of soft tissue properties under metatarsal heads…57
3.1.2 Development of tissue tester…………………………………………59
3.1.2.1 Multiple DOF foot positioning apparatus………………….59
3.1.2.2 Portable motorized indentor…………………………… …60
3.1.3 Assessing in-vivo tissue properties under metatarsal heads…… 62
3.1.3.1 Repeatability of MPEAI tests…………………………… 62
3.1.3.2 Force-displacement response of soft tissue under
metatarsal heads ……………………………………………… … 66
3.1.4 Derivation of material constants for plantar soft tissue…………….67
3.1.4.1 Classical Hertz contact theory………………………… ….67
3.1.4.2 Direct extraction of hyperelastic material constants… …67

3.1.4.3 Use of material constants in plane-strain FEM……….….72
3.2 Gait platform for measuring 3-D plantar stresses………………………… 75
3.2.1 3-D plantar stress measurement……………………………….……75
3.2.2 Design, fabrication and calibration of pressure/shear sensor… 77
3.2.2.1 Mechanical design of force sensor……………………….77
3.2.2.2 High-resolution sensor array………………………………78
3.2.2.3 Design of force sensor electronics……………………… 80
3.2.2.4 Sensor calibration………………………………………… 81
3.2.2.5 Construction of new gait platform…………………………84
3.2.3 Assessing sub-MTH loads………………………………………… 85
3.2.4 Calculation of shear traction ratio ………………………………….89

4
Chapter 4 Evaluation of model performance
4.1 Comparison of FEM results with experimental data…………………………93
4.1.1 Simulated muscular loads…………………………………………….94
4.1.2 Plantar stress distribution…………………………………………… 96
4.1.3 Metatarsal bone strain………………………………………………99
4.2 Sensitivity to flexor muscle force variations……………………………….…100
4.2.1 Role of gastrocnemius-soleus (G-S) muscle complex……… …100
4.2.2 Parametric study on F
AT
variations…………………………………104
4.2.2.1 Effects of Achilles tendon forces variations on movements
at the ankle and metatarsophlageal joints…………………………… 105
4.2.2.2 How Achilles tendon forces variations affects the forefoot
plantar pressure/shear stresses distribution………………………… 107
4.2.2.3 Effects of F
AT
variations on sub-MTH pressure peaks…110

4.3 Possible application of model in design of therapeutic footwear………….112
4.3.1 Insoles for stress relief under MTHs……………………………….112
4.3.2 Modification of insoles for foot-support interface applications… 114
4.3.3 Efficacy of modified insole in relieving loads under MTHs….… 115

Chapter 5 Conclusions
5.1 Summary……………………………………………………………… ……….118
5.2 Original Contributions…………………………………………………… … 120
5.3 Future Directions……………………………………………………………….121

Appendix A: Finite element analysis (FEA) …………………………………124
References…………………………………………………………………………133
Publications arising from the thesis………………………………………… 141




5
ABSTRACT
A musculoskeletal finite element model (FEM) of the human foot was developed
to investigate plantar stresses underneath metatarsal heads (MTHs) in the
forefoot. Vertical and shear forces at individual MTH sites were also measured in
vivo through use of a specially-designed novel gait platform system, which allows
three-dimensional interfacial stresses and local shear traction ratios to be
calculated. Average peak vertical pressures obtained are within the range of
published data based on commercial capacitance sensors. The measurements
from experiments were employed as inputs to the FEM model to define sliding
contact with Coulomb friction at the foot-support interface, as well as to validate
the foot model. The mechanical properties of the soft tissue under MTHs were
quantified by means of an instrumented indentation device. It was found that sub-

MTH pad tissue behavior is unique and dependent on metatarsophalangeal joint
dorsiflexion angle. The data obtained was also employed to determine the
material constants in the hyperelastic constitutive model adopted to describe
such tissue. This facilitated accurate estimation of the sub-MTH stresses induced
during foot-ground interactions.
Simulations were undertaken to predict changes in foot mechanism, in
terms of internal joint configurations and plantar loads distributions, which would
occur to accommodate any reduction in muscle effectiveness of the
gastrocnemius-soleus complex. The results correspond well with clinical
observations in diabetic patients who underwent tendo-Achilles lengthening
procedures. Stress reductions at individual MTHs were found to be site-specific

6
and possibly dependent on foot structures, such as intrinsic alignment of the
metatarsals. These highlight the clinical relevance of the model established in
terms of analyzing the foot mechanism. To illustrate possible application of the
model developed to therapeutic footwear intervention, the effects of modified
foot-support interfaces were also examined. This showed that inclusion of a
metatarsal support into a flat foam pad has the potential to relieve sub-MTH
plantar stresses. Forefoot plantar stress distribution is sensitive to positioning of
the metatarsal support, as well as the material used.
The level of complexity of the foot model established is unprecedented
and enables examination of the biomechanical interplay among muscular control,
bony joint movement, and foot/support interface interactions. This is also the first
attempt at such a comprehensive investigation of the foot mechanism. The
instrumented tissue tester and gait platform system developed are unique
experimental tools which facilitate determination of material model parameters,
as well as model validation. This is also a pioneer study of the efficacy of design
variables for therapeutic footwear in relieving contact stress, including plantar
shear, at sub-MTH region.

7
LIST OF TABLES
Tables Descriptions Page
1.2.2.1. A summary of plantar shear sensors reported in the literature
23
2.2.2.1.
Summary of FE model listing element type and material
properties for different model entities
37
2.2.4.2.
The material constants of the hyperfoam strain energy function
for the foam pad used as the foot-supporting interface
44
2.3.2.2.
The input forces in the muscles applied through the nodes
connected to tendon elements to drive the finite element foot
model
50
3.1.4.3.
The material properties used in the current finite element foot
model in comparison with those existing models for heel pad
71
3.2.2.1.

Calibration results of 9 strain-gauged sensors corresponding to
27 channels used in the 3 by 3 sensor array. For each channel
(AP: anterioposterior; ML: mediolateral; V: vertical) of an
individual sensor 6 measurements were performed.
Measurement errors were quantified as the standard deviation
of the maximum local force amplitude in the AP, ML, and vertical
directions, respectively. Average measurement errors are 2.8%,
3.1%, and 2.7% for directional forces in the vertical, AP and ML
axes, respectively
63
3.2.3.1.
Typical targeting errors showing the X
i
and Y
i
offset values for
each of the five successive trials at the 2nd MTH (PPP: peak
plantar pressure; PSS-AP: peak anterioposteiro shear stress;
PSS-ML: peak mediolateral shear stress)
87
4.1.2.1.
Summary of peak plantar stresses measured by sensor array
and predicted by the finite element model under individual
metatarsal heads (MTHs) (PSS-AP: peak anterioposteiro shear
stress; PSS-ML: peak mediolateral shear stress)
99

8
4.1.3.1.
The average strain in the 2nd metatarsal mid-shaft in

comparison with the cadaver study results by Sharkey et al.,
(1995)
99
4.2.2.1.
Variations of G-S complex muscle force (F
AT
) studied in the
present model sensitivity analysis
104
4.3.3.1.
Model predicted changes in plantar stress peaks under 2nd
MTH due to different foot support conditions. (MSP-5:
Metatarsal support positioned 5 mm proximal to the 2nd MTH)
116

9
LIST OF FIGURES
Figures Descriptions Page
1.2.1.1.
Diagram illustration of the effect of shear stress and pressure
on plantar tissue deformation modes underneath the metatarsal
head (MTH). The distortion associated with pure pressure is
compression. Shear stress results in tissue distortions in
parallel planes (i.e., angular deformation), which give rise to
elevated stress and tend to cause a “tear” within the tissue
close to bone prominences (Cavanagh and Ulbrecht, 2006)
20
2.1.2.1.
Segmentation of a human foot from individual coronal CT
slices. The bones were modeled as articulated parts enveloped
into a bulk soft tissue
31
2.1.2.2.
3-D solid model of foot geometry, including a bulk soft tissue
(A) and bones (B). Note that the foot skeletal was stabilized by
various ligaments
32
2.1.3.1.
Finite element mesh the foot (A) soft tissue, (B) internal bony
structures, and (C) various ligaments
33
2.2.4.1.

Stress-strain curve obtained from uni-axial compression test of
deformable four different types of foam padding materials
42
2.3.1.1.
Master/Slave approach to model contact interactions. The
contact behavior strictly follows the kinematic implications that
slave nodes cannot penetrate master surface segments
(Hibbert and Karssonn, 2006)
46
2.3.2.2.
Cut through the element mesh of the finite element model of
muscular foot and ankle complex, incorporating internal soft
tissue, skeletal structures, ligaments, plantar fascia, and
musculotendinous units for push-off simulation. Application of
muscle forces were simulated by force vectors align with the
tendons attached. F
AT
= Gastrocnemius-soleus complex, F
TIBP

= Tibialis posterior, F
FHL
= Flexor hallucis longus, F
FDL
= Flexor
49

10
digitorum longus, F
PB

= Peroneus brevis, and F
PL
= Peroneus
longus
2.3.3.1.
Relative sliding of points with contact constraint. Note the
possible evolution of contact between node 101 and its master
surface, BSURF, with sliding contact conditions with friction
52
2.3.4.1.
The finite element predicted stress response of a whole human
foot subjected to comprehensive musculoskeletal loading
corresponding to a heel-rise posture. Outcome measures of
primary interest in foot biomechanics, including stress
distributions of bulk soft tissue, metatarsal bones, ligaments
and plantar fascia, can be obtained
55
3.1.2.1.
Schematic diagram of the sub-Metatarsal Pad Elasticity
Acquisition Instrument (MPEAI), showing details of probe tip,
accommodation of sub-MTH pad and inside components of
actuator to drive probe tip
60
3.1.2.2.
(A) Photograph and (B) schematic diagram of test set-up,
showing ‘trapped’ 2nd sub-MTH pad and initial contact with
probe tip
61
3.1.2.3. (A) Displacement-time profile of indentation cycle for probe tip
61

3.1.3.1.
Typical tissue reaction force patterns obtained for 2nd sub-MTH
pad for similar indentation, but with MTP joint (MTPJ) at
different dorsiflexion angles. There was qualitative similarity in
all sets of curves for given testing conditions, and this indicated
that the experiments were well-controlled. All the curves
possess similar characteristic profiles, such as nonlinear
loading/unloading phases and a substantial force relaxation
component. Note that the sub-MTH pad tissue response is
dependent on MTP joint angle, i.e. lower stiffness at lower joint
dorsiflexion and increasing stiffness at higher joint angles
64
3.1.3.2.
Calculated initial stiffness (= peak tissue reaction force/ δ, δ1
mm)
65

11
3.1.3.3.
Calculated end-point stiffness (= peak tissue reaction force/
δ
max
)
65
3.1.3.4.
Force relaxation as percentages at δ
max
for 2nd Sub-MTH pad
of two normal subjects, for increasing MTP joint angles
65

3.1.4.1.
Values of material constants (

;

α) obtained through nonlinear
least squares (NLS) minimization using the deduced contact
equations for of the 1st-order Ogden hyperelastic model in
MATLAB. The calculated root-mean-square-error (RMSE) was
as small as 0.512 N with a R
2
value of 0.9821
70
3.1.4.2.
Comparison of the stress–strain behavior for the current sub-
MTH pad tissue model and those previously models: heel pad
by Erdemir et al. (2006), heel fat pad and skin model by Spears
et al. (2007)
71
3.1.4.3.
Application of the calculated material constants in a plane-strain
finite element model. The MTH geometry was simplified as a
sphere with a diameter of r. The r values was determined
through optimal surface fitting of the 3-D contour lines which
defines the actually shape of the 2nd MTH. The sub-MTH soft
tissue thickness t was defined as the shortest line distance
between the MTH and the plantar surface
72
3.2.2.1.
(A) Schematic diagram of sensor showing the positions of the

strain gauges and (B) photograph showing the attachment of
strain gauges to the front surfaces of the sensor body. Gauges
are also bonded to the rear surfaces. The positions of the
vertical and shear channels are also shown
77
3.2.2.2.
(A) Nine strain-gauged sensors were assembled into (B) a 3 by
3 sensor array; for a single strain-gauge channel, (C) the
electronic interface of a full-bridge circuit configuration was
shown
79




12
3.2.2.3.
A screen shot of the dynamic data acquisition (DAQ) system for
real-time strain measurements of the force sensor array using
LabVIEW (National Instrument). The block diagram of the
program is not shown
79
3.2.2.4. Calibration graphs showing linearity of the vertical channel
82
3.2.2.5. Calibration graphs showing linearity of the AP shear channel
82
3.2.2.6. Calibration graphs showing linearity of the ML shear channel
15
3.2.2.7.
(A) Schematic diagram of the gait platform with mounted sensor

array. (B) Photo of the foot making contact with the gait plate
that was embedded in a 7-meter walk way, sensor location in
relation to the placement of the foot plantar surface can be
monitored from the reflected mirror image by a high speed
camera (Fascam) positioned to capture the side view
84
3.2.3.1.
Method to create the “MTH template”: Images of the plantar
surface of the forefoot (A) without and (B) with sensor during
barefoot walking. Bony mark was used to identify the target
metatarsal site. (C) R1~R3 represents three non-collinear
markers for image registration between A and B. X
i
and Y
i
are
errors due to metatarsal site and sensor location discrepancy
86
3.2.3.2.
Force traces in vertical, AP and ML directions measured under
the 1st (A), 2nd (B), 3rd (C) and 4th (D) MTHs during barefoot
walking trails
88
3.2.3.1.
Local traction ratios obtained at sub-MTH area for the example
reading of horizontal (F
x
and F
y
) and vertical forces (F

z
) from
our customized pressure/shear sensor array
90
4.1.1.1.
Comparison of the muscular loads applied into the finite
element model with those calculated by Salathe and Arangio
(2002) and Gefen (2000)
95
4.1.2.1.
Model predicted (A) plantar pressure, (B) AP and (C) ML shear
distributions, and (D) VMS stress distributions in metatarsal
bones
97

13
4.2.2.1.
Correspondence of gastrocnemius-soleus muscle forces
(indicated as F
AT
) and the ankle () and metatarsophalangeal
() joints’ angles during heel rise. The amount of the muscle
forces in FAT was varied between 1620N (100% effectiveness)
and 972N (60% effectiveness). The angles were calculated
relative to the joint configurations at foot’s neutral position
105
4.2.2.2.
Percentage changes of local pressures under individual MTHs
for different ankle joint angles (). Nodal points just underneath
five MTHs were selected for analysis. Note the most significant

pressure reduction site was under the 2nd MTH for the
particular model analyzed in this study
107
4.2.2.3.
Changes in resultant plantar tissue shear stress distributions in
response to incrementally reduced gastrocnemius-soleus (G-S)
muscle forces.
108
4.2.2.4.
Percentage changes of local pressures under individual MTHs
for different ankle joint angles (). Nodal points just underneath
five MTHs were selected for analysis. Note the most significant
pressure reduction site was under the 2nd MTH for the
particular model analyzed in this study
110
4.3.2.1
Finite element model of a flat insole with an integrated
metatarsal support component interfacing with the forefoot
plantar surface
114
4.3.3.1.
Finite element predicted plantar anterio-posterior shear
stresses with different foot-supporting interfaces (A) rigid
ground, (B) soft foam pad, (C) a soft insole integrated with
metatarsal support
115
5.3.1.1.
FE mesh of a musculoskeletal foot model with flexors and
extensors for stance phase gait simulation
122

5.3.1.2.
A further gait platform system with a larger sensor array can be
installed in a typical gait lab for large-scale experiments
123


14
LIST OF ABBREVIATIONS AND SYMBOLS
a
Contact radius
AP Anterior-posterior
CT Computer tomography
DOF Degree-of-freedom
DAQ Data Acquisition
DM Diabetes mellitus
E
Young’s modulus
EVA Ethylene-vinyl acetate
F
AT
Gastrocnemius-soleus complex
F
TIBP
Tibialis posterior
F
FHL
Flexor hallucis longus
F
FDL
Flexor digitorum longus

F
PB
Peroneus brevis
F
PL
Peroneus longus
F
x

Anterior-posterior friction force
F
xy

Resultant shear forces
F
z

Normal force
FEM Finite element method
FHL Flexor hallucis longus
FDL Flexor digitorum longus
GRF Ground reaction force
G-S Gastrocnemius-soleus
LOPS Loss of protective sensation
MTPJ Metatarsophalangeal joint
MPEAI Metatarsal Pad Elasticity Acquisition Instrument
ML Medial-lateral
MRI Magnetic resonance imaging
MTHs Metatarsal heads
NLS Nonlinear least squares

15
PCSA Physiological cross-sectional areas
PB Peroneus brevis
PL Peroneus longus
PPP Peak plantar pressure
PSS-AP Peak anterior-posterior shear stress
PSS-ML Peak medial-lateral shear stress
R
Radius of the rigid indenter
TIBP Tibialis posterior
TMF Total muscle force
U
Strain-energy function
VMS Von Mises stresses
σ
Principal Cauchy stresses



Indentation pressure



Indentation strain
δ
max
Maximum probe depth

Loading/unloading rate





Ground state shear modulus
µ
Friction coefficient


Poisson’s ratio
λ
1
,

λ
2
, λ
3
Principal stretches








16

CHAPTER 1
INTRODUCTION





~If you want to accomplish the goals of your life, you have to
begin with the spirit~
Oprah Winfrey

17
1.1. Biomechanics of neuropathic diabetic foot ulcers
Diabetes mellitus (DM) is a syndrome of disordered metabolism characterized by
the common trait of excessive levels of glucose in the blood (e.g., when diabetes
is not under control). Sustained elevations of blood glucose are often not noted
by the patient; yet long term diabetes-associated lower extremity complications,
including distal peripheral neuropathy, to a lesser extent, peripheral vascular
disease, has been recognized for many years (Catterall et al., 1956). The
diabetic neuropathy leads to ‘loss of protective sensation’ (LOPS), whereby a
person is unable to feel minor trauma from repetitive mechanical stress during
daily activity that ultimately result in foot ulceration (Cavanagh et al., 1993). The
most costly and feared consequence is plantar ulceration that too often results in

partial or total amputation of the foot. The lifetime risk for a patient with diabetes
developing a foot ulcer could be as high as 25%, and diabetes related limb
amputation occurs 10 to 30 times more often in DM patients than in the general
population (Singh et al., 2005).
In the past, many in vivo and in vitro studies that addressed foot risks
have focused on delineating the potential overloading of the foot. Some early
investigators have utilized pressure transducers, and shown that sites of
maximum pressure coincided with plantar ulcer sites present in the patients with
neuropathic lesions (Stokes et al., 1975). The authors concluded that mechanical
loading acting on the foot plantar surface plays a “central pathogenic role” in the
onset of neuropathic ulcers. Since then, the predictive value of foot plantar
pressure (i.e. normal stress resulting from vertical force) has been widely

18
evaluated (Duckworth et al., 1985, Giacomozzi and Martelli, 2006, Stess et al.,
1997, Veves et al., 1992). Many investigators attempted to establish the
relationship between peak plantar pressures and the foot risks for tissue
breakdown. However, only moderate correlation between peak plantar pressure
and the occurrence of diabetic foot ulcers has been found so far (Armstrong et
al., 1998, Lavery et al., 2003). Recent studies have suggested that peak plantar
pressure may only be 65 % specific for the development of ulcers (Cavanagh et
al., 2000).
Many authorities thus believed that the shear component (i.e. resulting
from horizontal force) of the repetitive plantar loading, which has been largely
ignored previously, may be of equal importance to peak plantar pressure in
putting foot in greater risks for lesions (Delbridge et al., 1985). For plantar shear,
what we have observed is mainly based on force plate measurements during gait
analysis: the existence of medial-lateral (ML) and anterior-posterior (AP)
components of the ground reaction force (GRF), originating from foot-ground
interaction during walking. However, due to technical characteristics of force

plates, these only recorded the resultant GRF under the whole foot. Currently,
there is still a lack of suitable systems or approaches capable of characterizing
the three-dimensional stress states acting on localized anatomical sites such as
areas under metatarsal heads (MTHs), a common site for plantar ulcerations.
This chapter will provide a brief overview of the biomechanics of neuropathic
diabetic foot ulcers in general, and more attention will be given to emerging
techniques for shear measurement and modeling of foot-support interface.

19
1.2. Literature review
1.2.1. Plantar stresses and mechanical etiology of diabetic foot ulcers
A proper understanding of the etiology of foot ulcers may date back to the paper
by Catterall and Oakley et al., (1956) who drew attention to the peripheral
neuropathy as the most important contributory cause in the pathway to foot
ulcers. The diabetic neuropathy is an anatomically diffuse process that primarily
affects all three components of nerves (sensory, motor and autonomic), though
sensory symptoms are often most apparent. This leads to LOPS, whereby a
person is unable to feel minor trauma from repetitive mechanical stress during
daily activity that ultimately result in ulceration, which is now known as
neuropathic ulcers. Thus, in the causal pathway to neuropathic diabetic foot
ulcers, the major contributing factors are believed to be the presence of LOPS
due to peripheral neuropathy and repetitive excessive mechanical stresses
acting on the sole of the insensate foot during gait, such as walking.
Despite the fact that the plantar foot experiences three-dimensional force
vectors during gait, only the plantar pressure (arising from vertical component)
has been studied thoroughly. An early investigation conducted (Stokes et al.,
1975) using customized pressure transducers, was the first to show that sites of
maximum pressure coincided with plantar ulcer sites present in the patients with
neuropathic lesions. Since then, the predictive value of foot plantar pressure has
been widely evaluated (Duckworth et al., 1985, Giacomozzi and Martelli, 2006,

Stess et al., 1997, Veves et al., 1992). Many investigators attempted to establish
the relationship between peak plantar pressures and the foot risks for tissue

20
breakdown. However, only moderate correlation between peak plantar pressure
and the occurrence of diabetic foot ulcers has been found so far (Armstrong et
al., 1998, Lavery et al., 2003). Recent studies have suggested that peak plantar
pressure may only be 65 % specific for the development of ulcers (Cavanagh et
al., 2000). Many authorities thus believed that shear component (resulting from
horizontal force) of the repetitive loading, which has been largely ignored for
long, may be an equally important pathogenic factor to pressure in putting foot in
greater risks for lesions under local MTH sites (Fig. 1.2.1.1.).

MTH
Subcutaneous
tissue
Muscle
Fat
Skin
Pure Pressure
Shear + Pressure
Plantar Shear
MTH
Subcutaneous
tissue
Muscle
Fat
Skin
Pure Pressure
Shear + Pressure

Plantar Shear

Fig. 1.2.1.1. Diagram illustration of the effect of shear stress and pressure on plantar
tissue deformation modes underneath the metatarsal head (MTH). The distortion
associated with pure pressure is compression. Shear stress results in tissue distortions
in parallel planes (i.e., angular deformation), which give rise to elevated stress and tend
to cause a “tear” within the tissue close to bone prominences (Cavanagh and Ulbrecht,
2006).





21
1.2.2. Challenges in measuring plantar shear stresses
Because of the technical challenges for measuring the horizontal components,
anterior-posterior (AP) and medial-lateral (ML) shear, there is little known about
the actual distribution of the magnitude and direction of plantar shear during daily
activities, nor about the role that shear plays in causing plantar ulceration. So far,
only few devices have been designed to measure plantar shear stress during the
last decade. In 1980, Tappin et al. reported the first plantar shear sensor in the
literature. The device is based on magneto-resistive principle. Relative
displacements of two stainless steel discs induce measurable changes in
resistance from magneto-resistor which is proportional to the unidirectional shear
force. The apparent drawbacks of the work are 1) positioning and holding the
sensor in place could have had a perturbing effect on the local plantar stresses
(e.g., causing elevated forces) that were generated at contact interface; 2) it
measures shear in the longitudinal or AP direction only. Lord and colleagues
(1992) later enhanced this magneto-resistive device by allowing measurement to
be made in two orthogonal directions. By mounting the sensor into the ‘extra-

depth’ orthopaedic shoes, the authors measured shear stress of 20 to 60 kPa,
which are approximately 10% of the vertical stress values that are commonly
encountered. However, the device was only calibrated at quasi-static condition
and showed a large hysteresis of 18.5% of full-scale output with the range of 30N.
To avoid this problem, Akhlaghi and Pepper (1996) proposed a copolymer
piezoelectric film based sensor. This design is superior in linear response and
small hysteresis. Its disadvantages are that it is extremely sensitive to

22
temperature changes and thus not practical for clinical application where hostile
testing environments are expected. Davis and associates (1998, 2002) have
designed and tested a strain-gauged device that is composed of 16 sensors,
each 2.54ｘ2.54 cm
2
, capable of measuring three components of the applied
load. The device sensor showed excellent linearity (5%) and minimal hysteresis
(7.5%). Based on this design of the force sensor, the authors have made some
rudimentary measurements of planar shear stress in diabetic patients during
walking (Yavuz, M. 2007, 2008).
Apart from the foregoing review on shear sensors, Wang WC and Ledoux
WR (2005) presented a novel means of transducing plantar pressure and shear
stress with a fiber-optic sensor. They constructed a map of normal and shear
stresses based on observed macrobending through the intensity attenuation from
the physical deformation of two adjacent perpendicular fibers. To overcome the
hurdle in calibrating sensors at the individual level, Mackey and Davis (2006)
developed a prototype system based on optical methods, where each sensing
element is made from the same polycarbonate material. Therefore, the authors
claim that sensing linkages may be calibrated only once under zero-load
conditions over the entire array of sensors. Although each sensor has clear
indication of both shear and pressure and a good spatial resolution of 8 mm, it is

still subject to technology limitations such as some hysteresis in the system and
slightly larger distance between stress sensor linkages. A summary of these
reported plantar shear sensors with their technical specifications are given in
Table 1.2.2.1.

23
Table 1.2.2.1. A summary of plantar shear sensors reported in the literature
Author Year Sensor Type Resolution Capacity
Tappin et al.,
(1980)
1980 magneto-resistive
16 mm in
diameter
Longitudinal
(AP) shear
Lord et al., (1992) 1992 magneto-resistive
15.96 mm in
diameter in-
shoe use
Longitudinal
(AP) shear
Akhlaghi&Pepper,
(1996)
1996
copolymer
piezoelectric film
10 x 10 mm
Bi-axial shear
(AP plus ML)
Davis et al., (1998) 1998

Strain gage
technology
2.5 x 2.5 cm
2

Pressure and
shear sensor
Wang et al., (2005) 2005 fiber-optic sensor 1 × 1 cm
2

Pressure and
shear sensor
Mackey and Davis,
(2006)
2006
Polycarbonate
stress linkage
(optical based)
8 mm in
diameter
Pressure and
shear sensor