METHODS IN THE MEASUREMENT OF BIOMECHANICAL RESPONSE TO
MODIFICATIONS IN FOOTWEAR DESIGN
By
Jerrod E. Braman

A THESIS
Submitted to
Michigan State University
in partial fulfillment of the requirements
for the degree of
MASTER OF SCIENCE
Engineering Mechanics
2010

ABSTRACT
METHODS FOR MEASURING BIOMECHANICAL CHANGES DUE TO
MODIFICATIONS IN FOOTWEAR MIDSOLE CHARACTERISTICS
By
Jerrod E. Braman
Orthopaedic problems can be caused or accentuated by both normal and abnormal gait
characteristics.

Traditionally, footwear modifications and orthotic interventions have been

limited to internal and external wedges and other assorted orthotics. This thesis will discuss
various methods to measure biomechanical responses of human subjects in response to changes
in footwear design. Chapter 1 provides an overview of what is currently known regarding
interventions in relation to providing medial compartment knee osteoarthritis relief as well as
limiting excessive rear-foot motion. Chapter 2 discusses a method for measuring the center of
pressure in multiple steps using an insole pressure measurement system and evaluating the effect
that varying midsole stiffness has on the center of pressure. Chapter 3 presents a dynamic
segmental model for knee adduction moment measurements and how a full lateral stiffened shoe
and five other prototype shoes with varying midsole stiffness change that moment in healthy
subjects. Chapter 4 presents a method for measuring rear-foot motion and evaluates a new
footwear technology designed to reduce excessive rear-foot motion. Chapter 5 is a summation of
the thesis and recommendations for future studies. The information in this thesis could be
helpful in providing strategies for changing location and stiffness values of footwear midsoles to
elicit positive biomechanical responses in human subjects during gait.

ACKNOWLEDGMENTS
I would first like to thank Dr. Roger C. Haut for direction, support, and patience through the past
five years. I would not be where I am now without his help. I would also like to thank Dr.
Dianne Ulibarri for guiding me to the field of biomechanics. This research would not be
possible without the technical help from Clifford Beckett. In addition, I would like to thank Dr.
John Powell for his guidance and Dr. Tamara Reid-Bush for sitting on my committee. I also
want to thank my wife, Leah, for her support and being so understanding and patient throughout
this whole process. My family also deserves thanks, as they have always given their support
through the good times and the tough times. In addition, I would like to thank Wolverine World
Wide, Inc. for providing the footwear used for this research. Last but not least, I would like to
thank all of my friends and fellow students that have helped me through this journey.

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TABLE OF CONTENTS
LIST OF TABLES............................................................................................................v
LIST OF FIGURES ........................................................................................................ vi
CHAPTER 1
INTRODUCTION ............................................................................................................1
REFERENCES ...............................................................................................................15
CHAPTER 2
CENTER OF PRESSURE AND PRESSURE DISTRIBUTION PROFILES OF
SUBJECTS WEARING VARIABLE MIDSOLE DENSITY FOOTWEAR ..........18
REFERENCES ...............................................................................................................40
CHAPTER 3
DOES THE KNEE ADDUCTION MOMENT CHANGE WHEN WEARING
PRODUCTION MODEL FULL LATERAL STIFFENED AND PARTIAL
LATERAL WITH STIFFENED MEDIAL HEEL SHOES?....................................42
REFERENCES ...............................................................................................................80
CHAPTER 4
EVALUATION OF A NOVEL FOOTWEAR DESIGN AND ITS EFFECTS ON REARFOOT MOTION AND MECHANICAL STIFFNESS COMPARISON TO
COMMERICAL MOTION CONTROL SHOES.....................................................82
REFERENCES .............................................................................................................102
CHAPTER 5
SUMMARY AND RECOMMENDATIONS FOR FUTURE STUDY.......................103
REFERENCES .............................................................................................................108
APPENDIX A:
APPENDIX B:
APPENDIX C:
APPENDIX D:
APPENDIX E:

DATA FROM CHAPTER 2 ..............................................................110
DATA FROM CHAPTER 3 ..............................................................142
BODYBUILDER CODE FOR CHAPTER 3 ....................................151
BODYBUILDER CODE FOR CHAPTER 4 ....................................164
OTHER BIOMECHANICAL PROJECTS ........................................167

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LIST OF TABLES
Table 2.1 – Subject Demographics .................................................................................23
Table 3.1 – Marker Locations and Names for Static and Dynamic Trials .....................59
Table 3.2 – Subject Demographics .................................................................................66
Table 3.3 – Average Knee Adduction Moment by Subject ............................................72
Table 3.4 – Average Overall Knee Adduction Moment .................................................74
Table 3.5 – Gendered Average of the Knee Adduction Moment ...................................75
Table 4.1 – Summary of Angle Change from Braking to First Peak..............................95
Table 4.2 – Summary of Angle Change from First Peak to Midstance..........................96
Table 4.3 – Summary of Angle Change from Braking to Midstance .............................97
Table 4.4 – Stiffness Values for B06-21 Insole with Red Wave Disk ...........................98
Table 4.5 – Stiffness Values for Select Motion Control Shoes ......................................98
Table B.1 – Knee Moment Data for Subject F1 ...........................................................143
Table B.2 – Knee Moment Data for Subject F2 ...........................................................144
Table B.3 – Knee Moment Data for Subject F3 ...........................................................145
Table B.4 – Knee Moment Data for Subject F4 ...........................................................146
Table B.5 – Knee Moment Data for Subject M1 ..........................................................147
Table B.6 – Knee Moment Data for Subject M2 ..........................................................148
Table B.7 – Knee Moment Data for Subject M3 ..........................................................149
Table B.8 – Knee Moment Data for Subject M4 ..........................................................150

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LIST OF FIGURES
Figure 1.1—Illustration of the orientation of a knee adduction moment. View is of a left
leg in a posterior orientation. GRF stands for ground reaction force and Madd
stands for the adduction moment. ...................................................................3
Figure 1.2—Yasuda and Sasaki diagram for spatial changes in the calcaneus leading to an
angular decrease in the extended mechanical axis due to orthotic use ...........5
Figure 2.1—The Parotec System and a sample insole....................................................24
Figure 2.2—Picture of the experimental shoe based on the Merrell Full Pursuit
Model ............................................................................................................25
Figure 2.3—Top view of the location and Asker C values for the experimental shoes.
Right shoe configurations are shown – left shoe configurations mirror their
counterparts...................................................................................................26
Figure 2.4—Definitions of terms and visualizations for reading Parotec data...............29
Figure 2.5—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 2 ..............32
Figure 2.6—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 3 ..............33
Figure 2.7—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 4 ..............34
Figure 2.8—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 5 ..............35
Figure 2.9—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 6 ..............36
Figure 3.1—Diagram of the calculation of the knee adduction moment using the FRFV
method...........................................................................................................45
Figure 3.2—Anatomical Model ......................................................................................47
Figure 3.3—Massless Segmental Model ........................................................................47
Figure 3.4—Quasi-Static Segmental Model ...................................................................48
Figure 3.5—Full Dynamic Model...................................................................................48
Figure 3.6—Comparison of FRFV and segmental models from one trial of one
subject ...........................................................................................................58
Figure 3.7—Anterior, Lateral, and Posterior Views of the Anatomical Model with Marker
Set .................................................................................................................60
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Figure 3.8—Anatomical skeletal landmark references for the marker set .....................61
Figure 3.9—Photo of the Merrell Full Pursuit shoe style for the neutral control shoe and
full lateral stiffened shoe (top) and the location and Asker C Stiffness values
for the midsole configurations (bottom) .......................................................68
Figure 3.10—Photo of the Merrell Full Pursuit shoe style for the five T-Form shoes (top)
and the location and Asker C Stiffness values for the midsole configurations
(bottom).........................................................................................................69
Figure 3.11—Subject walking in a T-Form model shoe across the force plate..............70
Figure 3.12—A typical knee adduction moment vs. time plot .......................................71
Figure 4.1—Anterior view of supination (a) and pronation (b)......................................82
Figure 4.2—Posterior view of the modified insert with a Wave Disk (left) and the control
insert (right) ..................................................................................................84
Figure 4.3—Top view of a red Wave Disk.....................................................................84
Figure 4.4—Bottom view of a blue Wave Disk .............................................................85
Figure 4.5—Bottom view of a right WDCT insole with Wave Disk placed in the antipronation setting............................................................................................85
Figure 4.6—Posterior view of the red Wave Disks in the anti-pronation setting...........86
Figure 4.7—Posterior view of the red Wave Disks in the anti-supination setting .........86
Figure 4.8—Bates B06-21 Boots....................................................................................88
Figure 4.9—Marker set for the posterior calcaneus........................................................89
Figure 4.10—Marker set for the dorsal forefoot.............................................................89
Figure 4.11—Picture of the posterior calcaneus marker set placed in the flexible receiving
piece ..............................................................................................................90
Figure 4.12—From top left, clockwise—(a) Right foot taped with flexible receiving pieces
exposed; (b) Right foot with marker sets attached; (c) Right foot with boot and
marker sets attached......................................................................................91
Figure 4.13—Graphical Explanation of Standard Gait Events.......................................92

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Figure 4.14—Location of Instron stiffness testing using 19mm x 42 mm rectangular
indenter head.................................................................................................94
Figure A.1 – Definitions of terms and visualizations for reading Parotec data............111
Figure A.2 – Subject M1 Data – Shoe 1 vs. Shoe 2 .....................................................112
Figure A.3 – Subject M1 Data – Shoe 1 vs. Shoe 3 .....................................................113
Figure A.4 – Subject M1 Data – Shoe 1 vs. Shoe 4 .....................................................114
Figure A.5 – Subject M1 Data – Shoe 1 vs. Shoe 5 .....................................................115
Figure A.6 – Subject M1 Data – Shoe 1 vs. Shoe 6 .....................................................116
Figure A.7 – Subject M2 Data – Shoe 1 vs. Shoe 2 .....................................................117
Figure A.8 – Subject M2 Data – Shoe 1 vs. Shoe 3 .....................................................118
Figure A.9 – Subject M2 Data – Shoe 1 vs. Shoe 4 .....................................................119
Figure A.10 – Subject M2 Data – Shoe 1 vs. Shoe 5 ...................................................120
Figure A.11 – Subject M2 Data – Shoe 1 vs. Shoe 6 ...................................................121
Figure A.12 – Subject M3 Data – Shoe 1 vs. Shoe 2 ...................................................122
Figure A.13 – Subject M3 Data – Shoe 1 vs. Shoe 3 ...................................................123
Figure A.14 – Subject M3 Data – Shoe 1 vs. Shoe 4 ...................................................124
Figure A.15 – Subject M3 Data – Shoe 1 vs. Shoe 5 ...................................................125
Figure A.16 – Subject M3 Data – Shoe 1 vs. Shoe 6 ...................................................126
Figure A.17 – Subject F1 Data – Shoe 1 vs. Shoe 2.....................................................127
Figure A.18 – Subject F1 Data – Shoe 1 vs. Shoe 3.....................................................128
Figure A.19 – Subject F1 Data – Shoe 1 vs. Shoe 4.....................................................129
Figure A.20 – Subject F1 Data – Shoe 1 vs. Shoe 5.....................................................130
Figure A.21 – Subject F1 Data – Shoe 1 vs. Shoe 6.....................................................131
Figure A.22 – Subject F2 Data – Shoe 1 vs. Shoe 2.....................................................132

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Figure A.23 – Subject F2 Data – Shoe 1 vs. Shoe 3.....................................................133
Figure A.24 – Subject F2 Data – Shoe 1 vs. Shoe 4.....................................................134
Figure A.25 – Subject F2 Data – Shoe 1 vs. Shoe 5.....................................................135
Figure A.26 – Subject F2 Data – Shoe 1 vs. Shoe 6.....................................................136
Figure A.27 – Subject F3 Data – Shoe 1 vs. Shoe 2.....................................................137
Figure A.28 – Subject F3 Data – Shoe 1 vs. Shoe 3.....................................................138
Figure A.29 – Subject F3 Data – Shoe 1 vs. Shoe 4.....................................................139
Figure A.30 – Subject F3 Data – Shoe 1 vs. Shoe 5.....................................................140
Figure A.31 – Subject F3 Data – Shoe 1 vs. Shoe 6.....................................................141

ix

CHAPTER 1 – INTRODUCTION
Gait is the pattern of movement of the limbs during locomotion. Orthopaedic problems
can be caused or accentuated due to both abnormal and normal gait. The knee is an especially
susceptible joint where injury and other orthopaedic problems can manifest due to gait
characteristics. Specifically, the knee joint is susceptible to degeneration of cartilage, commonly
referred to as osteoarthritis. Other lower limb problems have been linked to the problem of
excessive pronation, a common gait abnormality. Traditionally, footwear modifications have
been limited to internal and external wedges and other assorted orthotics. This introduction will
discuss methods that have been used to delay the onset or slow the progression of osteoarthritis
by attempting to modify knee adduction moments. There will also be a discussion of some of the
issues related to excessive foot pronation and the methods used in an attempt to limit such
rotations.
Knee Joint
Osteoarthritis of the knee joint can be caused or accelerated by gait. According to Mow
and Huiskes, osteoarthritis is characterized by a breakdown of articular cartilage (Mow and
Huiskes 2005). The knee can be divided into medial and lateral compartments, which are on the
inside and outside of the knee, respectively. Knee osteoarthritis occurs in about 10% of adults
after the age of 55 (Davis 1988).

Within this group, osteoarthritis occurs in the medial

compartment almost 10 times as frequent as in the lateral compartment (Ahlback 1968).

One

mechanism that is associated with the onset or progression of osteoarthritis is that of an
abnormally large adduction moment (Fisher, Dyrby et al. 2007). Approximately 60% to 80% of
the load transmitted to the knee goes through the medial compartment and is largely due to this
moment (Crenshaw, Pollo et al. 2000). It has also been suggested that there is a correlation

1

between the intensity of the knee adduction moment and the severity of medial compartment
disease (Ueda, Dyrby et al. 2003). The severity of osteoarthritis can be measured several
different ways, but the Kellgren-Lawrence scale is the most common method. The KellgrenLawrence scale has five separate descriptive categories of a joint based on radiological features
that help quantify the severity of osteoarthritis. These five radiological features are listed below
(Kellgren and Lawrence 1957).

1) The formation of osteophytes on the joint margins or, in the case of the knee joint, on
tibial spines.
2) Periarticular ossicles; these are found largely in relation to the distal and proximal
interphalangeal joints.
3) Narrowing of joint cartilage associated with sclerosis of subchondral bone.
4) Small pseudocystic areas with sclerotic walls situated usually in the subchondral
bone.
5) Altered shape of the bone ends, particularly in the head of the femur.

The five categories of osteoarthritis based on these features are as follows: (0) None; (1)
Doubtful; (2) Minimal; (3) Moderate; (4) Severe. A radiologist would classify an individual into
one of these five categories based on an x-ray of the affected area.
A knee adduction moment is centered at the knee joint and oriented in the coronal plane.
This concept is illustrated in Figure 1.1. There have been numerous studies to investigate the

2

efficacy of different adduction moment reducing techniques for either placating existing medial
compartment osteoarthritis or delaying its onset. These methods include surgical and nonsurgical interventions.

The surgical methods include a high-tibial osteotomy or total knee

replacement. The high-tibial osteotomy is an invasive surgery in which a wedge of bone is taken
out of the tibia in order to redistribute the forces through the knee. Due to the nature of the
surgery, a long recovery period ensues following this operation. Bracing devices, such as canes
or crutches, may be necessary for up to 10 weeks post-surgery. This bracing is in addition to the
rehabilitation that is necessary. A total knee replacement is used in cases where there is little or
no cartilage left on the end of the bones.

GRF

Madd

Figure 1.1 – Illustration of the orientation of a knee adduction moment. View
is of a left leg in a posterior orientation. GRF stands for ground reaction force
and Madd stands for the adduction moment.

3

Less invasive methods have been proposed, primarily in the last 20 years, to provide an
alternative to the surgical approaches such as a high-tibial osteotomy or knee replacement. The
use of orthotics and other non-surgical interventions have been shown to decrease the knee
adduction moments. For example, Yasuda and Sasaki measured the change in position of the
center of gravity between barefoot and with a wooden lateral wedge with an incline of 5 degrees
placed under the entire foot (Yasuda and Sasaki 1987).

The findings of this study suggest a

more conservative method of decreasing static medial knee compressive loads. The explanation
for this effect was that the calcaneus becomes more valgus when the lateral wedge is used. This
is in addition to a more vertical angle of the lower limb mechanical axis. This change in
mechanical axis would cause a decrease in the adduction moment at the knee due to a shorter
moment arm between the knee joint center and the ground reactive force vector. Figure 1.2
describes this change in spatial relation of the “extended mechanical axis of the lower limb” with
and without the use of the lateral wedge. The extended mechanical axis is the line drawn from
the center of the femoral head to the point of heel strike contact on the calcaneus. It should be
noted that the angle between the femur and the tibia and fibula complex does not change – only
the angle of the calcaneus becomes more valgus to impart the change in angle.
Yasuda and Sasaki (1987) also noted that a change in the tibiofemoral angle does not
need to occur in order for a decrease in adduction moment. The other suggested method for a
reduction in adduction moment occurs with a lateral shift in the origin of the ground reaction
force, causing the force to pass closer to the center of the knee joint.

4

γ

γ

Figure 1.2 – Yasuda and Sasaki diagram for spatial changes in the calcaneus leading to
an angular decrease in the extended mechanical axis due to orthotic use
There have been several studies trying to quantify the effect of the lateral wedge orthotic.
The experiments have largely been designed to quantify the benefits of the lateral wedge orthotic
for people who already have knee osteoarthritis (Kerrigan, Lelas et al. 2002). Yet, there have
also been studies on people who do not have any symptoms of knee osteoarthritis (Crenshaw,
Pollo et al. 2000; Kakihana, Akai et al. 2004; Fisher, Dyrby et al. 2007). There has also been
reseach investigating the role of orthotics during running and other sporting activities in healthy
subjects (Nigg, Nurse et al. 1999; Mundermann, Nigg et al. 2003; Nigg, Stergiou et al. 2003) or
injured subjects (Kakihana, Torii et al. 2005). Kakihana et. al. (2005) “assessed the biomechanic
effects of wearing a lateral wedge on the subtalar joint movement during gait in athletes with and
without an unstable lateral ankle.” The focus of the study was to find the joint kinematic effects
of a 6 degree full lateral wedge with the stable and unstable ankle. The authors document a
decrease in knee adduction moment in the 50 male athletes studied, regardless of their subtalar

5

stability status. In addition, they also made note of a lateral shift in the center of pressure. It was
noted that the origin of the ground reaction force vector was shifted lateral when using the 6
degree full lateral wedge as compared to a zero degree platform during dynamic walking trials.
This lateral center of pressure shift coupled with a decrease in adduction moment is important to
note because it supports the method of adduction moment reduction put forth by Yasuda and
Sasaki (1987).
Nigg et. al. also performed experiments to find kinematic, center of pressure, and leg
joint moment responses to flat, half- and full-medial and half- and full-lateral orthotics (Nigg,
Stergiou et al. 2003).

Data were collected from 15 healthy male subjects with no knee

osteoarthritis. For the lateral wedge condition, the lateral height of the orthotic was 4.5 mm
higher than the medial border. The opposite was true for the medial wedge. The authors
document that the full lateral orthotics generated a statistically significant lateral shift in the
center of pressure but not a consistent change in the adduction moment at the knee. In addition,
no other insert condition significantly and consistently changed the center of pressure. The only
significant finding with the full medial insole was an increase in the knee external rotation
moment. The COP finding is consistent with Kakihana et al. (2004), but the authors suggested
that the lateral COP shift using partial length orthotics was not highly correlated to a reduction in
the knee adduction moment. They also state that the overall reactions to orthotics are often
subject-specific effects and cannot be related to the general population.
The effectiveness of lateral wedged insoles has also been evaluated in people with knee
osteoarthritis. One group tested the biomechanical efficacy of 5 and 10 degree lateral wedges in
patients with at least a category 3 knee on the Kellgren-Lawrence scale (Kerrigan, Lelas et al.
2002). The authors document that the 10 degree lateral wedge decreases the peak adduction

6

moment by 8%, while the 5 degree lateral wedge decreases the peak moment by 6%. In addition
to these measurable changes, a comfort questionnaire was also given to the subjects. They found
that even though the 10 degree lateral wedge was more effective at reducing peak knee adduction
moment, the insert also provided a significant amount of discomfort. In essence, the subjects
were trading discomfort at the knee for discomfort in the shoe. If the patient is unwilling to wear
the orthotic, the benefit of decreasing the adduction moment is lost. This implies that in order
for an orthotic intervention to be successful, it also must be comfortable.
In addition to studies being performed on athletes and persons with osteoarthritis, studies
quantifying the effect of a lateral wedge in the general population have been documented in
current literature. Crenshaw et al (2000) studied the effects of a lateral wedged insole in 17
healthy subjects. The study used a 5 degree full lateral orthotic in a velocity-controlled walking
experiment. The results of their study suggest no change in joint angles at the ankle, knee, or hip
and also no change in kinetics at the hip or ankle. The study documents a significant reduction
in the external adduction moment at the knee joint. This study is important because it was the
first to document the reduction of knee adduction moment in a healthy, non-athletic population
with the use of a full lateral wedge orthotic.
Following Crenshaw et al (2000), Kakihana et al (2004) performed a study on healthy
adults using full lateral wedged insoles. The authors’ aim was to try and find a correlation
between the decrease in knee adduction moment and the change in moment at the subtalar joint.
The authors used 3 and 6 degree lateral wedged insoles on 10 healthy adults (5 male, 5 female).
They document that the larger, 6 degree wedged insole intervention significantly decreases the
knee adduction moment. The authors indicate that the decrease in knee adduction moment was
most likely due to a lateral shift in the center of pressure. The lateral shift in pressure was found

7

by measuring the increase in the length of the moment arm at the subtalar joint. From this
finding, it was assumed that the more lateral location of the center of pressure caused a smaller
distance of the moment arm at the knee joint in the adduction/abduction plane and subsequent
reduction of the knee adduction moment.
Fisher et al (2007) also addressed the issue of reducing the knee adduction moment in
healthy subjects with the use of an in-shoe intervention. In addition to the traditional lateral
wedge, the authors proposed the use of a dual-density mid-sole with the lateral component of the
mid-sole being stiffer than the medial half of the insole. A total of 14 subjects (9 male and 5
female) were used in the study. The intervention conditions tested in the study were a 4 degree
full lateral wedge orthotic, an 8 degree full lateral wedge orthotic, a 120% lateral stiffness shoe
(a shoe that is 20% stiffer in the lateral half than the medial half), and a 150% lateral stiffness
shoe (a shoe that is 50% stiffer in the lateral half than the medial half). These four intervention
conditions were compared to the personal walking shoes of each study participant. The authors
report that there was a significant decrease in the knee adduction moment when using the 150%
stiffness shoes, as well as the 4 and 8 degree full lateral wedge orthotics. The investigators also
indicated that the 120% stiffness shoes show a statistical trend towards decreasing the moment.
When the methods of reduction were compared, they document that the 8 degree full lateral
wedge provides the largest reduction in adduction moment. As seen in a previous study, full
lateral wedges with high degree values provide the most reduction in adduction moment, but
they can also be uncomfortable. While the 150% stiffness shoe did not provide as much relief as
the two wedged conditions, it may be a more effective way to decrease the knee adduction
moment in patients without osteoarthritis because of the minimal impact on the subject. This use

8

of variable density mid-sole characteristics is intriguing, but the tests in this study did not use
production model shoes.
A second study was performed by this group using variable density midsole technology
using subjects with medial compartment knee osteoarthritis (Erhart, Mundermann et al. 2008).
Using a constant-stiffness shoe and a variable-stiffness intervention shoe, 79 subjects were tested
at three different paces of walking gait – slow, self-selected, and fast. The variable stiffness shoe
had a lateral half that was 1.3-1.5 times stiffer than the medial half. The Asker C values – a
measure of material stiffness – for the medial half were 55+2, and the lateral half had values
between 70-76+2. Subjects who were bilaterally affected with osteoarthritis were evaluated on
the limb with the most pain, as subjectively determined by the subject. An inverse dynamics
model was used with a six marker set and a force plate in order to determine kinetic and
kinematic parameters. It was shown that the peak knee adduction moment was reduced at all
speeds of walking, and the largest reduction in moment was found with an increase in gait
velocity.
Lower Limb Problems Caused by Excessive Pronation
There are many possible problems throughout the body that originate at or around the
ankle joint during gait. The focus of the studies here will be to discuss the role of excessive
pronation as it relates to injury mechanisms. Excessive foot pronation has been proposed to
increase the possibility for patellofemoral pain syndromes, shin splints, Achilles tendinitis,
plantar fasciitis, and stress fractures (Nigg 2001). For purposes of this thesis, pronation and
eversion motions were considered to be the same. In actuality, pronation is a complicated triplanar motion that includes motion of the individual bones in the foot, while eversion is often

9

considered to be only the rotation of the foot about its long axis. There have been several
strategies proposed to change the degree of foot pronation.
One excessive pronation reduction method is taping. The low-dye arch support is one
taping technique that is employed by athletic trainers as a method to reduce the amount of foot
pronation (Schulthies and Draper 1995). Another commonly used method to control inversion or
eversion of the foot is by the use of posting an insole or creating a custom-molded insole, making
it biomechanically difficult for the wearer to pronate or supinate past a certain point of rotation.
These in-shoe modifications are described in many articles (Stacoff, Reinschmidt et al. 2000;
Mundermann, Nigg et al. 2003; Nigg, Stergiou et al. 2003; Kakihana, Akai et al. 2005;
Kakihana, Torii et al. 2005).
In the Stacoff, et. al. (2000) study, five male subjects were outfitted with intracortical
bone pins and tested in running sandals with different medial foot orthoses in an attempt to
accurately measure skeletal motion during running. Reflective marker triads were attached to the
pins for motion analysis as subjects ran across a forceplate. The three conditions tested were a
neutral orthotic, a posterior medial orthotic, and an anterior medial orthotic. The posterior
orthotic was designed to support the foot from the calcaneus to the medial malleolus, while the
anterior orthotic supported the arch region of the foot. In each condition, support was created by
posting or by a raised area in the orthotic. There was no difference in the reduction of eversion
between the two insole conditions or between the posterior orthotic, the anterior orthotic, and the
control insole. However, four of the five subjects did show a decrease in maximum foot eversion
between 1 and 3 degrees, but because of only having a small subject pool the difference was not
statistical.

10

Mundermann et. al. (2003) performed a study also using running sandals, but not bone
pin arrays. Instead, 21 subjects (12 female, 9 male) had markers attached to the skin surrounding
the calcaneus. Kinematic data were collected using a motion capture system and evaluated using
a software program developed by the University of Calgary. Four different insole conditions
were tested – a control insole, a full medial posted insole, a neutral custom-made shell, and a
custom shell with full medial posting. They document that posting alone significantly reduces
the amount of maximum eversion, while molding and posting and molding changes the amount
of tibia rotation, but not the degree of foot eversion.
In addition to measuring moment and center of pressure data, the 2003 Nigg et. al. study
evaluated 15 healthy male subjects while running in five different shoe configurations in terms of
their capability to change the amount of foot inversion and eversion. One shoe insert was the
unaltered insert that arrived with the shoe. The test conditions were full lateral, full medial, half
lateral, and half medial posting. The half inserts were placed at the metatarsal region. Posting
was 4.5 mm either medial or lateral. Kinetic and kinematic data were captured using a motion
capture system and a force plate. An inverse dynamics method was used to calculate resultant
joint moments. It was found that the full lateral insert condition increased the amount of initial
foot inversion (1.8 + 2.1 degrees), and the full medial insert condition decreased the maximum
eversion position (-1.5 + 1.3 degrees). In addition, the total eversion motion was larger in the
full lateral (2.1 + 1.7 degrees) and smaller in the full medial condition (-2.0 + 1.5 degrees).
The first of two Kakihana et. al. studies (2004) assessed the kinetics and kinematics of 10
subjects (5 men, 5 women) while they walked in three different insole conditions – no wedge, a
3-degree lateral angle wedge, and a 6-degree lateral angle wedge. This study was unable to find

11

a significant difference in the average pronation angle during the stance phase when comparing
the 6-degree wedge with the no wedge condition.
The second Kakihana et. al. study (2005) was conducted with 50 males in the no wedge
and 6-degree lateral wedge condition. In the subject population, 25 subjects were considered to
have an unstable lateral ankle while the other 25 subjects were considered to be healthy controls.
Once again, the amount of rotation was not significantly different between subject groups or
between insole conditions. What the researchers did find, however, was that the subtalar joint
moment in the frontal plane was different between the two different insole conditions for both
subject populations.
Most of the studies mentioned above have had similar experimental designs and have
used similar equipment. In general, a motion analysis system coupled with a force plate is used
to properly calculate joint kinetics and kinematics. The use of retro-reflective marker sets
coupled with extensive knowledge of normal body motion allows an investigator to determine
and define motion of body segments with a high degree of accuracy. It has been shown that a
full inverse dynamics approach is necessary to accurately calculate moments about joint centers
(Winter 2005). Most studies have been performed using male rather than female subjects, or at
least a majority of male subjects. Men and women differ anatomically, especially in the lower
limbs. In general, women have a larger quadriceps angle (Q-angle) than men. This change in
the mechanical axis of the lower limb, coupled with foot shape, could lead to a difference in the
response to in-shoe interventions. It has also been documented that “female feet and legs are not
simply scaled-down versions of male feet but rather differ in a number of shape characteristics,
particularly the arch, the lateral side of the foot, the first toe, and the ball of the foot”
(Wunderlich and Cavanagh 2001). With these concepts in mind, thesis objectives were set.

12

Thesis Objectives
The objectives of this thesis were threefold. The first objective was to determine if
variable midsole density shoes could alter the foot center of pressure and pressure distribution
patterns of healthy male and female subjects during gait. The second objective of this thesis was
to determine if the knee adduction moment changes when wearing production model full lateral
stiffened and experimental shoes with a stiffened medial heel and a changing location of a pocket
of lateral stiffness in healthy subjects. The third objective of the thesis was to determine if a
novel user-controlled footwear technology could alter the amount of rear-foot motion in healthy
subjects.

13

REFERENCES

14

References
Ahlback, S. (1968). "Osteoarthrosis of the knee. A radiographic investigation." Acta Radiol
Diagn (Stockh): Suppl 277:7-72.

Crenshaw, S. J., F. E. Pollo, et al. (2000). "Effects of lateral-wedged insoles on kinetics at the
knee." Clin Orthop Relat Res(375): 185-92.

Davis, M. A. (1988). "Epidemiology of osteoarthritis." Clin Geriatr Med 4(2): 241-55.

Erhart, J. C., A. Mundermann, et al. (2008). "A variable-stiffness shoe lowers the knee adduction
moment in subjects with symptoms of medial compartment knee osteoarthritis." Journal of
Biomechanics 41: 2720-2725.

Fisher, D. S., C. O. Dyrby, et al. (2007). "In healthy subjects without knee osteoarthritis, the
peak knee adduction moment influences the acute effect of shoe interventions designed to reduce
medial compartment knee load." J Orthop Res 25(4): 540-6.

Kakihana, W., M. Akai, et al. (2005). "Effects of laterally wedged insoles on knee and subtalar
joint moments." Arch Phys Med Rehabil 86(7): 1465-71.

Kakihana, W., M. Akai, et al. (2004). "Changes of joint moments in the gait of normal subjects
wearing laterally wedged insoles." Am J Phys Med Rehabil 83(4): 273-8.

Kakihana, W., S. Torii, et al. (2005). "Effect of a lateral wedge on joint moments during gait in
subjects with recurrent ankle sprain." Am J Phys Med Rehabil 84(11): 858-64.

Kellgren, J. H. and J. S. Lawrence (1957). "Radiological assessment of osteo-arthrosis." Ann
Rheum Dis 16(4): 494-502.

Kerrigan, D. C., J. L. Lelas, et al. (2002). "Effectiveness of a lateral-wedge insole on knee varus
torque in patients with knee osteoarthritis." Arch Phys Med Rehabil 83(7): 889-93.

15

Mow, V. C. and R. Huiskes (2005). Basic Orthopaedic Biomechanics and Mechano-Biology.
Philadelphia, Lippincott Williams & Wilkins.

Mundermann, A., B. M. Nigg, et al. (2003). "Foot orthotics affect lower extremity kinematics
and kinetics during running." Clin Biomech (Bristol, Avon) 18(3): 254-62.

Nigg, B. M. (2001). "The role of impact forces and foot pronation: a new paradigm." Clin J Sport
Med 11(1): 2-9.

Nigg, B. M., M. A. Nurse, et al. (1999). "Shoe inserts and orthotics for sport and physical
activities." Med Sci Sports Exerc 31(7 Suppl): S421-8.

Nigg, B. M., P. Stergiou, et al. (2003). "Effect of shoe inserts on kinematics, center of pressure,
and leg joint moments during running." Med Sci Sports Exerc 35(2): 314-9.

Schulthies, S. and D. O. Draper (1995). "A Modified Low-Dye Taping Technique to Support the
Medial Longitudinal Arch and Reduce Excessive Pronation." Journal of Athletic Training 30(3):
3.

Stacoff, A., C. Reinschmidt, et al. (2000). "Effects of foot orthoses on skeletal motion during
running." Clin Biomech (Bristol, Avon) 15(1): 54-64.

Ueda, S., C. O. Dyrby, et al. (2003). Adduction Moment, Radiographic Disease Severity, and
Pain in Osteoarthritis of the Knee. 2003 Summer Bioengineering Conference. Key Biscayne,
Florida.

Winter, D. A. (2005). Biomechanics and Motor Control of Human Movement. Hoboken, John
Wiley & Sons, Inc.

Wunderlich, R. E. and P. R. Cavanagh (2001). "Gender differences in adult foot shape:
implications for shoe design." Med Sci Sports Exerc 33(4): 605-11.

16

Yasuda, K. and T. Sasaki (1987). "The mechanics of treatment of the osteoarthritic knee with a
wedged insole." Clin Orthop Relat Res(215): 162-72.

17

CHAPTER 2 – CENTER OF PRESSURE AND PRESSURE DISTRIBUTION PROFILES
OF SUBJECTS WEARING VARIABLE MIDSOLE DENSITY FOOTWEAR

Introduction
The study of the use of in-shoe interventions such as orthotics or wedges to elicit a
positive biomechanical response (e.g. gait correction) has its roots in basic studies starting in the
late 20th century. Sasaki and Yasuda studied orthotics in the 1980’s through the use of wedges to
elicit a change in the gait of patients with problems in their lower limbs (Yasuda & Sasaki,
1987). The novel idea that a non-invasive alternative to surgery, a custom fit insert, could
alleviate or mitigate joint pain and disease has become a routine method of treatment for many
lower limb ailments. The static studies performed by Yasuda and Sasaki led to a number of
questions related to the efficacy of this kind of cost effective intervention. With the advent of the
digital age and better measurement equipment, researchers have taken on the task of thoroughly
evaluating the performance of footwear interventions in people with and without lower limb
problems such as osteoarthritis, excessive joint motion, and general pain.
It has been shown that the changes in the center of pressure can be related to changes in
joint moments (Fuller, 1999; Kakihana, Akai et al., 2005; Xu, Akai, Kakurai, Yokota, & Kaneko,
1999). It has been suggested that these changes in joint moments could alleviate some lower
limb problems (Crenshaw, Pollo, & Calton, 2000; Kerrigan et al., 2002; Yasuda & Sasaki, 1987).
The center of pressure (COP) is defined as the location at which the resultant of the ground
reaction forces can be statically placed at any point in time (Winter, 2005). The foot COP curve
is a track showing the time history of the COP during a step. Several different types of foot
orthotics have been shown to change the COP using different measurement tools. Force plates
are sometimes used to measure the center of pressure (Kakihana, Akai, Yamasaki, Takashima, &

18

Nakazawa, 2004; Xu et al., 1999). The drawback to force plates is that these researchers are
limited in the number of concurrent steps that can be studied using force plates. For example, if
a laboratory employs two force plates, they are limited to two center of pressure plots per trial.
In order to record multiple COP plots, a different device must be used. One such set of devices
that can measure the COP are plantar pressure measurement systems.
In a 1992 paper by Rose, et. al., an FScan Pressure System was used to measure the
effects of external heel wedges and their effects on plantar pressures and the center of pressure
(Rose, Feiwell, & Cracchiolo, 1992). Eleven subjects (6 men, 5 women; age range 25-45 years)
were tested in five conditions – no heel wedge, quarter-inch lateral heel wedge, quarter-inch
medial heel wedge, half-inch lateral heel wedge, and a half-inch medial heel wedge. The authors
found that the quarter-inch heel wedge did not significantly change the pressure distribution,
while the half-inch lateral wedge decreased pressures under the third, fourth and fifth metatarsal
heads, while increasing the pressure underneath the first and second metatarsal heads.

In

addition, the half-inch medial wedge significantly decreased the pressure under the first and
second metatarsals but did not significantly increase the pressure under the third, fourth, and fifth
metatarsals. The COP curves were shifted laterally when using the quarter- and half-inch medial
heel wedge, while the shift was medial during the use of the two lateral wedge designs. The
half-inch varieties of each wedge shifted the COP more than the quarter-inch wedge. The largest
shifts were found in the midfoot and metatarsal regions, with little to no COP change in the heel
area.
Xu, et. al. describe changes in the COP when using Thomas and reverse Thomas heel
modifications in addition to a rocker bar that was placed in three different locations (Xu et al.,
1999). A Thomas heel is an external medial heel wedge, and a reverse Thomas heel is an

19

external lateral heel wedge. Twenty subjects (10 male, 10 female) participated in this study. It
was found that when wearing a shoe with the Thomas heel, the COP was shifted towards the
lateral border of the shoe when compared to the standard heel and the reverse Thomas heel
shifted the COP towards the medial border of the shoe as compared to the standard heel. The
COP lines were measured with a force plate and changes in the COP were made by finding the
area between the COP line and the center line of the foot.
Shoe orthotics can also be placed in between the foot and the shoe – not just on the
outside of the shoe. A 2003 Nigg article discussed the effect that different in-shoe orthotics had
on the COP as compared to a neutral condition using a PEDAR plantar measurement system to
measure the COP (Nigg et al., 2003). The researchers compared a neutral shoe condition to four
different inserts – a medial forefoot wedge, lateral forefoot wedge, full length medial wedge, and
a full length lateral wedge. Each posting was 4.5mm at its maximum height. The authors noted
that the only consistent and statistical change in the COP occurred with the full lateral wedge.
The shift of the COP was towards the lateral border with a full lateral wedge – the opposite effect
of an externally applied lateral wedge. A separate group of researchers also found changes in the
COP towards the side of posting.
A 2004 paper by Van Gheluwe and Danaberg also used a plantar pressure measurement
system to measure changes in COP when using seven different medial and lateral heel and
forefoot wedges (Van Gheluwe & Dananberg, 2004). Twenty-three subjects were tested using
the Footscan measurement system to obtain COP and plantar pressure distribution data. The
forefoot inserts were 3 degrees valgus, 3 degrees varus, and 6 degrees varus. The heel inserts
were 4 degrees valgus, 4 degrees varus and 8 degrees varus. Valgus wedges are the same as
lateral wedges, while varus wedges are analogous to medial wedges. These inserts were all

20

compared to a neutral condition. It was noted that the varus (medial) wedge caused a medial
shift in the area that it was designed to affect (forefoot or heel), while the valgus (lateral) wedge
caused a lateral shift the same area. It was noted that the COP did not change in the region
opposite where the wedge was located. This was also true for changes in pressure distribution.
That is, the forefoot wedge conditions did not change the heel pressure distribution and the heel
wedge conditions did not change the forefoot pressure distributions.
One drawback of using an internal or external foot orthotic is the potential for user
discomfort. A research study noted that some patients would remove and discontinue the use of
a wedge because of this discomfort, even knowing their pain may return (Kerrigan et al., 2002).
This potential disadvantage of orthotics has led researchers to explore other options to change the
lower body joint moments. One such alternative to a wedge is a variable stiffness midsole. To
date, little is known about the effects on the COP when changing patterns and densities of
midsole materials and if it has the same effect on COP as traditional orthotics.
For this study, the Parotec System, an in-shoe pressure measurement device, was used to
collect plantar pressure data over a given time period. From this plantar pressure data, COP plots
could be created. In order for the Parotec to effectively create the center of pressure plots, it
needed to be accurate and its measures repeatable. A 2000 study by Chesnin et. al. took on the
task of validating the Parotec in-shoe pressure measurement system by comparing the COP
generated by the Parotec to the COP from an AMTI force plate. The study indicated that the
Parotec COP data were highly correlated with the COP data generated by the force plate
(Chesnin, Selby-Silverstein, & Besser, 2000). This result allowed for the use of the Parotec
System to draw accurate conclusions about possible changes in the COP. Therefore, this section

21

of the thesis will discuss the methods associated with an attempt to change the center of pressure
by changing the midsole density characteristics of a production model shoe.
The various midsole configurations used in the current study were designed by Wolverine
World Wide, Inc. for the purpose of altering the foot center of pressure lines and pressure
profiles. One objective of their pattern designs was to try and keep the COP along the centerline
of the shoe. In order to test these designs in production-type shoes, an in-shoe plantar pressure
measurement system was used because of its ability to capture multiple steps as well as for its
ease of tracking the COP line for comparison purposes.

Materials and Methods
Subjects
Six subjects (3 men and 3 women) were recruited from a local running club. The
participants had no prior history of injury or surgery to the lower body and ran a minimum of 10
miles per week. The average age was 54.0 years for the men and 29.7 years for the women. The
average height and weight for the men were 172.0 cm and 68.4 kg respectively, while the
women’s average height and weight were 161.7 cm and 55.3 kg respectively.

22

Table 2.1 – Subject Demographics
For interpretation of the references to color in this and all other figures, the reader is referred to
the electronic version of this thesis
Subject Demographics
Subject Code Height (m) Weight (kg)
M1
1.68
61.4
M2
1.73
62.5
M3
1.75
81.4
F1
1.63
56.8
F2
1.65
59.1
F3
1.57
50.0

Plantar Pressure Measurement System
The pressure profile and center of pressure data were collected using the Parotec
System® plantar pressure measuring device (Paromed, Neubeurn, Germany; Figure 2.1). In this
study, the pressure profile was determined as the average pressure of each sensor during each
individual step. The Parotec insoles contain 24 silicon-filled bladders spaced across the plantar
surface with each bladder containing a piezoresistive sensor. Data were collected at 120 Hz
using a belt-mounted controller and recorded on a PCMCIA memory card. Following each test
sequence the data were transferred to a personal computer and analyzed with Parotec system
software version 4.02.

23

23 24

22
21

20

18 19

17

15 16

13 14

11

7

12

8

10

9
6

5
3
1

4
2

Figure 2.1 – The Parotec System and a sample insole
Footwear
The shoes were specially designed for these tests, but they were production-quality
footwear based on the Merrell Full Pursuit trail running shoe (Figure 2.2). Six pairs of US men’s
size 9 shoes and six pairs of US women’s size 8 shoes were used in the study. Each pair of shoes
had a different midsole configuration (i.e. the stiffness of the midsole varied at different locations
across the plantar surface of the shoe). Each midsole had varying densities of ethylene vinyl
acetate (EVA). The Asker C scale was used to determine the hardness of the EVA material in
each midsole region. The control shoe had a uniform EVA density across the entire midsole
with an Asker C value of 53 (Figure 2.3). Larger Asker C values indicate stiffer materials.

24

Figure 2.2 – Picture of the experimental shoe based on the Merrell Full Pursuit model

25

53

53

53

53

53

73

53

53

53

53

73

73

Shoe 1

Shoe 2

Shoe 3

43 63

53

53

73

43

73

53

53

73

73

73

Shoe4

Shoe 5

Shoe 6

Figure 2.3 – Top view of the location and Asker C values for the experimental
shoes. Right shoe configurations are shown – left shoe configurations mirror
their counterparts.

26

Experimental Procedure
Prior to all tests, the factory insole was removed from the shoe. Sensor insoles matching
the size of the shoes were inserted into each pair of test shoes. The factory insoles were removed
in order to try and maximize the potential stiffening characteristics of each shoe by having the
subject’s foot as close to the midsole as possible. The Parotec insoles were then connected to the
controller by the use of lead wires. These wires were held to the subject’s ankles with Velcro
straps so that the small data boxes on the insoles were not subjected to excessive movement and
subsequently damaged. The insoles were zeroed as the subject raised their feet off the ground.
Data were recorded during three trials of running over a distance of approximately ten foot
strikes for both the left foot and the right foot. The subject was told to run at a comfortable, selfselected pace. Velocity was not recorded, as a comfortable pace for one subject may not have
been comfortable for another and the goal of the study was to find out what happens at a
comfortable running pace. The running surface was a tiled floor. Data collection was started
and stopped by remote control. Data were collected after constant gait velocity was reached and
prior to deceleration during each trial.

Data Analysis
The following parameters were measured with the Parotec Software: peak pressure in
each sensor, the distribution of pressures, and the trace of the center of pressure. The peak
pressure was the maximum pressure generated at each sensor of the Parotec insole during a trial.
The pressure distribution allowed for a visual evaluation of high and low pressure regions across
the plantar surface for shoes with different midsole density patterns.

27

When viewing Parotec data, a green box indicates a higher pressure in the second shoe
compared to the first shoe (e.g. Shoe 1 vs Shoe 2 green box means a higher pressure in Shoe 2
compared to Shoe 1 in that particular sensor). A red COP line indicates the control shoe, while
the blue COP line indicates the comparison shoe. The dominant foot, as informed by the subject,
was analyzed in the study.

Due to the way the pressure analysis software was set up,

comparisons could only be made between the same number steps across different trials. For
example, the center of pressure line from the control shoe step 1 could only be compared to step
1 of the other shoes. Figure 2.4 shows a typical step of comparison Parotec data. The steps that
are presented in Figures 2.5 through 2.9 were picked because they best represented the trends
from each subject. Results are presented as comparisons between a test shoe and the control
shoe.

28

Red COP line
for the control
shoe

Blue COP line for
the comparison
shoe

A green box indicates
higher pressure in
comparison shoe than
control shoe. A larger
box means a greater
difference.

A red box indicates
lower
pressure
in
comparison shoe than
control shoe. A larger
box means a greater
difference.

Figure 2.4 – Definitions of terms and visualizations for reading Parotec data

29

Results
Shoe 1 vs. Shoe 2 (Figure 2.5)
No significant differences were observed between the COP lines for any subject. It was
noticed that all three female subjects had higher maximum pressures on the lateral border of the
mid-foot and slightly lower pressures on the medial border in Shoe 2 as compared to Shoe 1 in a
majority of step comparisons. This does not necessarily correlate to a COP change because the
pressure values compared are maximum pressure values during a step and are independent of
time. There was also a decrease in pressure at the distal phalange of the first metatarsal for all
six subjects.
Shoe 1 vs. Shoe 3 (Figure 2.6)
Four subjects (F1, F2, F3, and M1) exhibited a lateral shift in the COP lines for a
majority of the steps in Shoe 3 as compared to the control shoe (Shoe 1). These shifts occur in
the midfoot region between the heel and metatarsals. There were no consistent areas of changes
in plantar pressure between Shoe 3 and the control shoe for the subjects tested.
Shoe 1 vs. Shoe 4 (Figure 2.7)
None of the subjects exhibited a consistent change in the COP between the test shoe
(Shoe 4) and the control shoe (Shoe 1). It was noticed that there was an increase in maximum
st

pressure at the base and a decrease in maximum pressure at the head of the 1 metatarsal in a
majority of the steps of each trial for five subjects (three women, two men). It is also noted that
the toe-off of the COP occurs earlier in the foot in shoe 4 as compared to the control shoe for all
steps of all six subjects. This phenomenon is not noticed in any other shoe comparison.
Shoe 1 vs. Shoe 5 (Figure 2.8)

30

Three subjects (F1, F2, and F3) exhibited a consistent shift in COP to the lateral border of
the shoe in Shoe 5 as compared to Shoe 1. Subject M1 had several trials of a lateral shift in
COP, but it was not a change that was seen in a majority of the steps analyzed. Four subjects
(F1, F2, F3, and M1) had a region of higher maximum pressure along the lateral border of the
middle half of the foot in Shoe 5 as compared to Shoe 1.
Shoe 1 vs. Shoe 6 (Figure 2.9)
There was no consistent change in COP from Shoe 6 as compared to Shoe 1 for any of
st

the six subjects. Four subjects (F1, F2, F3, and M3) had lower pressures in the tip of the 1
metatarsal in a majority of steps in each trial in Shoe 6 as compared to Shoe 1.

31

Subject F1

Subject F2

Subject F3

Shoe 1 vs. Shoe 2

53

53

53
53

Shoe 1
Subject M1

53

Subject M2

Subject M3

53

53
73
Shoe 2

Figure 2.5—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 2

32

Shoe 1 vs. Shoe 3

53

Subject F2

Subject F1

Subject F3

53

53
53

Shoe 1

73

53

Subject M1

Subject M2

Subject M3

53
73
Shoe 3

Figure 2.6—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 3

33

Shoe 1 vs. Shoe 4
Subject F1

53

Subject F2

Subject F3

53

53
53

Shoe 1

43

63

Subject M1

Subject M2

Subject M3

53
73

Shoe 4

Figure 2.7—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 4

34

Shoe 1 vs. Shoe 5
Subject F2

Subject F1

53

Subject F3

53

53
53

Shoe 1

53

73

Subject M1

Subject M2

Subject M3

53
73
Shoe 5

Figure 2.8—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 5

35

Shoe 1 vs. Shoe 6
Subject F2

Subject F1

53

Subject F3

53

53
53

Shoe 1

43

73

Subject M1

Subject M2

Subject M3

53
73

Shoe 6

Figure 2.9—Comparison of Control Shoe (Shoe 1) and Experimental Shoe 6

36

Discussion
The current study focused on changes in the foot pressure distribution and the center of
pressure between a neutral midsole shoe condition and several shoes with unique midsole density
characteristics. It was hypothesized that the shoes would behave like an internal wedge and shift
the COP toward the wedged components as documented in previous literature (Kakihana, Akai et
al., 2005; Nigg et al., 2003; Van Gheluwe & Dananberg, 2004). However, the two variable
midsole designs that exhibited a shift in the COP had a common area of increased medial heel
stiffness causing a lateral shift in the center of pressure (Shoes 3 and 5). This result is similar to
the literature regarding external heel wedges (Rose et al., 1992; Xu et al., 1999). It was noticed
that changes in the COP occurred in the midfoot area and not at the heel or forefoot regions of
the foot. This is important to note because Shoes 3 and 5 have exactly the opposite forefoot
composition – Shoe 3 had a stiff medial forefoot and Shoe 5 had a stiff lateral component. In
addition, the shoe with only medial heel stiffening did not create changes in the center of
pressure location. These differences suggest that changes in midsole density may not affect the
center of pressure in the forefoot region, but it is required to have different stiffness values in the
forefoot region to elicit a response.
While there have been several studies that have incorporated the use of variable midsole
density shoes (Erhart, Mundermann, Elspas, Giori, & Andriacchi, 2008; Fisher, Dyrby,
Mundermann, Morag, & Andriacchi, 2007), this study was one of the first to evaluate such a
wide distribution of different midsole EVA densities and their effects on the center of pressure.
It was also a unique study in that it used production-model shoes and not just modified footwear.
The rather unique distributions of high and low stiffness regions were supplied by Wolverine
World Wide, Incorporated personnel, with no particular reason the designs shown. A full lateral

37

stiffened shoe was not provided for this study because it was not considered to be a shoe model
that could ever be sold. A full lateral stiffened shoe could possibly increase ankle pronation
which was considered to go against running shoe manufacturing protocol according to Wolverine
World Wide, Inc.
It should be noted that while differences in pressure and center of pressure lines were
recorded, they were not necessary large in magnitude. This result coincides with previous
literature that quantified the amount of change in the COP (Kakihana, Torii et al., 2005; Nigg et
al., 2003; Rose et al., 1992). This may be a positive trait, however, as large changes in pressure
distribution or medial and lateral COP tracks may lead to foot instability and possible injury
(Fuller, 1999). This is a very important aspect that needs to be kept in mind when evaluating
shoes for potential commercial use. One interesting finding from this study was that even in the
shoes where shifts in the center of pressure were observed, the movement did not necessarily
manifest itself in every step.

This result is similar to the 2003 Nigg et. al. study which

documented changes in the center of pressure not occurring in every step (Nigg et al., 2003). In
addition, the two midsole configurations that shifted the COP (Shoes 3 and 5) only affected the
COP in the midfoot region and not at heel strike or toe-off. This is important to consider if the
biomechanical change desired manifests itself in these two regions.
These data could be useful to improve current orthotic prescription. If a patient is not
comfortable with the footwear modification, they are less likely to continue to use it. In this
study, an alternative to an orthotic wedge was introduced as a way to make the footwear
intervention more comfortable while being able to control the pressure distribution and center of
pressure. Further development and refinement of the locations and stiffness values of the
midsole density pattern may prove to be a useful tool when orthotic wedges are not appropriate.

38

REFERENCES

39

References
Chesnin, K. J., Selby-Silverstein, L., & Besser, M. P. (2000). Comparison of an in-shoe pressure
measurement device to a force plate: concurrent validity of center of pressure
measurements. Gait and Posture, 12, 6.
Crenshaw, S. J., Pollo, F. E., & Calton, E. F. (2000). Effects of lateral-wedged insoles on kinetics
at the knee. Clin Orthop Relat Res(375), 185-192.
Erhart, J. C., Mundermann, A., Elspas, B., Giori, N. J., & Andriacchi, T. P. (2008). A variablestiffness shoe lowers the knee adduction moment in subjects with symptoms of medial
compartment knee osteoarthritis. Journal of Biomechanics, 41, 2720-2725.
Fisher, D. S., Dyrby, C. O., Mundermann, A., Morag, E., & Andriacchi, T. P. (2007). In healthy
subjects without knee osteoarthritis, the peak knee adduction moment influences the
acute effect of shoe interventions designed to reduce medial compartment knee load. J
Orthop Res, 25(4), 540-546.
Fuller, E. A. (1999). Center of pressure and its theoretical relationship to foot pathology. J Am
Podiatr Med Assoc, 89(6), 278-291.
Kakihana, W., Akai, M., Nakazawa, K., Takashima, T., Naito, K., & Torii, S. (2005). Effects of
laterally wedged insoles on knee and subtalar joint moments. Arch Phys Med Rehabil,
86(7), 1465-1471.
Kakihana, W., Akai, M., Yamasaki, N., Takashima, T., & Nakazawa, K. (2004). Changes of
joint moments in the gait of normal subjects wearing laterally wedged insoles. Am J Phys
Med Rehabil, 83(4), 273-278.
Kakihana, W., Torii, S., Akai, M., Nakazawa, K., Fukano, M., & Naito, K. (2005). Effect of a
lateral wedge on joint moments during gait in subjects with recurrent ankle sprain. Am J
Phys Med Rehabil, 84(11), 858-864.
Kerrigan, D. C., Lelas, J. L., Goggins, J., Merriman, G. J., Kaplan, R. J., & Felson, D. T. (2002).
Effectiveness of a lateral-wedge insole on knee varus torque in patients with knee
osteoarthritis. Arch Phys Med Rehabil, 83(7), 889-893.
Nigg, B. M., Stergiou, P., Cole, G., Stefanyshyn, D., Mundermann, A., & Humble, N. (2003).
Effect of shoe inserts on kinematics, center of pressure, and leg joint moments during
running. Med Sci Sports Exerc, 35(2), 314-319.

40

Rose, N. E., Feiwell, L. A., & Cracchiolo, A., 3rd. (1992). A method for measuring foot
pressures using a high resolution, computerized insole sensor: the effect of heel wedges
on plantar pressure distribution and center of force. Foot Ankle, 13(5), 263-270.
Van Gheluwe, B., & Dananberg, H. J. (2004). Changes in plantar foot pressure with in-shoe
varus or valgus wedging. J Am Podiatr Med Assoc, 94(1), 1-11.
Winter, D. A. (2005). Biomechanics and Motor Control of Human Movement (3rd ed.).
Hoboken: John Wiley & Sons, Inc.
Xu, H., Akai, M., Kakurai, S., Yokota, K., & Kaneko, H. (1999). Effect of shoe modifications on
center of pressure and in-shoe plantar pressures. Am J Phys Med Rehabil, 78(6), 516-524.
Yasuda, K., & Sasaki, T. (1987). The mechanics of treatment of the osteoarthritic knee with a
wedged insole. Clin Orthop Relat Res(215), 162-172.

41

CHAPTER 3 – DOES THE KNEE ADDUCTION MOMENT CHANGE WHEN
WEARING LATERALLY STIFFENED PRODUCTION MODEL SHOES?
Introduction
Osteoarthritis is the breakdown of articular cartilage in a joint (Mow & Huiskes, 2005)
and occurs in the knee joint of approximately 10% of all people age 55 and older (Davis, 1988).
People with osteoarthritis of the knee are 10 times more likely to have the degenerative disease
in the medial compartment as compared to the lateral compartment (Ahlback, 1968). In addition,
women are more likely than men to develop osteoarthritis in the knee joint (Felson et al., 1997).
An increase in the knee adduction moment can accentuate, accelerate, and even induce medial
compartment knee osteoarthritis (Fisher, Dyrby, Mundermann, Morag, & Andriacchi, 2007).
The knee adduction moment has also been linked to being responsible for transmitting 60% to
80% of the total load that occurs in the knee (Crenshaw, Pollo, & Calton, 2000). There are
several invasive and non-invasive methods that have been found to decrease knee adduction
moments in people with and without medial compartment knee osteoarthritis.
One surgical method that has been shown to decrease the knee adduction moment in
people with medial compartment osteoarthritis is a high tibial osteotomy. Prodromos et. al.
found that the knee joint adduction moment decreased when a wedge of bone was taken out of
the tibia, thereby redistributing the forces in the knee joint (Prodromos, Andriacchi, & Galante,
1985). Non-invasive techniques have also been shown to be effective at reducing the adduction
moment in symptomatic osteoarthritis patients. Kerrigan et. al. found that 5 and 10 degree
wedges were effective at reducing the adduction moment (Kerrigan et al., 2002), while Kakihana
et. al. found that 6 degree wedges also significantly reduced the adduction moment in patients
with knee joint osteoarthritis (Kakihana, Akai et al., 2005). Kuroyanagi et. al. added subtalar
strapping to a 6 degree lateral wedge to also significantly decrease the knee adduction moment

42

(Kuroyanagi et al., 2007). Lateral wedges have also been shown effective at reducing the
adduction moment in non-symptomatic people.

Crenshaw et. al., Kakihana et. al., and

Fisher et. al. all found that lateral wedged insoles decrease the adduction moment in healthy
subjects (Crenshaw et al., 2000; Fisher et al., 2007; Kakihana, Akai, Yamasaki, Takashima, &
Nakazawa, 2004). In addition, Fisher et. al. also proposed a new, non-invasive method for
decreasing the adduction moment at the knee joint by the use of a dual-density midsole. They
found that by increasing the stiffness of the lateral half of an outsole/midsole combination by
20% and 50% more than the medial half, the knee adduction moment was significantly decreased
at a rate similar to a 4 degree lateral wedge (Fisher et al., 2007). One drawback to this study was
that the lateral stiffened shoes were prototype shoes that were not based on a conventional shoe
build.
With this knowledge, this thesis chapter set out to accomplish four objectives. The first
objective of this study was to develop a biomechanical model to determine the knee joint
adduction moment. The second objective was to determine if a full, lateral stiffened midsole in a
production shoe would decrease knee joint adduction moments. The third objective of the study
was to determine if shoes with a stiff medial heel component and a changing location of a pocket
of lateral stiffness could decrease the knee adduction moment. The fourth objective of the study
was to determine if the response to the change in midsole stiffening was different between male
and female subjects.
Biomechanical Model
Because it is difficult and expensive to surgically implant force transducers in the joints
of subjects, different techniques must be used to measure the desired effect.

In order to

determine the knee joint adduction moment during walking gait, a biomechanical model was

43

developed. Two methods for modeling joint moments are the floor reaction force vector (FRFV)
method and the inverse dynamics solution in a segmented body model. Both model types require
the use of some kind of system that can give the position of joint centers and the relative location
of the ground reaction force, but the segmental model requires more information to compute
joint moments.
The floor reaction force vector method is one way to calculate moment data. It calculates
the moment by finding the product of the ground reaction force and the perpendicular distance
from the joint center (d) in the different planes of motion. An example of how this method can
be used to calculate the knee adduction moment is shown in Figure 3.1. The force components
in the frontal plane (F) are multiplied by the perpendicular distance from the knee joint center (d)
to obtain the adduction moment.

44

Posterior View

Knee Joint
Center
Force (F)
Vector

Z-Axis

d

Vertical
to ground
Y-Axis
(Perpendicular
to the Track)

Force Vector
Origination
(COP)

θ

Force Plate

Figure 3.1 – Diagram of the calculation of the knee adduction moment using the FRFV method

There are significant drawbacks to this method, however. Winter outlined these problems in the
third edition of Biomechanics and Motor Control of Human Movement (Winter, 2005). The first
drawback to the method is that the moment values do not account for the acceleration of the
segments below the joint in question.

For example, using the FRFV method to measure

moments at the knee do not take into account the mass and acceleration of the foot and shank
segments. According to Wells, the error is negligible at the ankle, small but significant at the
knee, and quite significant at the hip (Wells, 1981). A second problem is that the method only
shows the cause of the moment and not what the body is actually doing. The example Winter

45

gives is with respect to the ankle joint. When the ankle is performing a plantarflexion moment,
the ground reaction force is forced forward. If the ankle muscles were not firing, the FRFV
would stay under the ankle joint. With this knowledge, the FRFV was not chosen as the method
to measure the knee adduction moment.
The second type of model used to compute joint moments in the body is the segmental
model.

Segmental models are characterized by breaking down the human anatomy into

simplified segments. The first level of a segmental model only takes into account that each body
segment acts on its own with respect to the ground reaction force and is known as a massless
model. The second level of a segmental model includes the effect of the weight of each segment.
This is known as a quasi-static model. The top level of a segmental model includes everything
from the massless and quasi-static models and then adds the effect of the accelerated moments of
inertia from each segment and is called a full dynamic model. The different segmental model
types can be seen in Figures 3.2 through 3.5.

46

Figure 3.2 -- Anatomical
Model

Hip Joint Center

Knee Joint Center

Ankle Joint Center

Figure 3.3 --Massless
Segmental Model

47

Hip Joint Center
Thigh Mass
Knee Joint Center

Shank Mass
Ankle Joint Center
Foot Mass
Figure 3.4 --Quasi-Static
Segmental Model

Hip Joint Center
Thigh Mass, Moment of Inertia
Knee Joint Center

Shank Mass, Moment of Inertia
Ankle Joint Center
Foot Mass, Moment of Inertia
Figure 3.5 --Full Dynamic
Model

48

In the current studies, anthropometric data from Dempster (1955) was used to determine the
location of the mass centers for the various body segments and for each segment weight for use
in the dynamic model. Dempster performed experiments on cadavers to establish locations and
segment weights as a function of total body mass. His data is still used in most inverse dynamic
biomechanical models. Sample calculations comparing the FRFV, the massless segmental, the
quasi-static segmental, and the full dynamic model can be found following the marker set
description.
Model Terms and Sample Calculations
The following section details the terms and equations used to determine the knee
adduction moment using the FRFV method and the three segmental models. All four models
needed the medial/lateral and vertical components of the reaction force from the force plate in
addition to the medial/lateral origin of the force vector, also known as the center of pressure
(COPy). Also, the four methods all needed the location of the knee joint center, while three
segmental methods also needed the ankle joint center location. These values, in addition to the
mass of the foot and shank segments and the spatial location of these masses, were known
quantities either from measurement or definition. The mass and location of the foot and shank
segments were calculated based on work by Dempster (1955), the force and center of pressure
data were measured by the force plate, and the knee and ankle joint centers were located using
the motion capture system.

49

FRFV Method
Using trigonometry, the perpendicular distance from the knee joint center to the
projection of the force vector in the medial/lateral and vertical plane (distance d) was calculated
for the FRFV method. This value was not required for the segmental models. As can be seen in
Figure 3.6, the resultant knee adduction moment is different for the FRFV method compared to
the Full Dynamic Segmental model. This is due to the fact that the FRFV model does not use
segments and does not take into account for the inertial properties of the foot and shank
segments.

FRFV Method – Terms
Kadd – Knee Adduction Moment
Fy-z – Projection of force vector in the Y-Z plane (medial/lateral and vertical force)
Fy – Force component in the Y – direction (medial/lateral)
Fz – Force component in the Z – direction (vertical)
d – Perpendicular distance from knee joint center to Fy-z
KJCy – Y component of the location of the knee joint center
KJCz – Z component of the location of the knee joint center
COPy – Y component of the center of pressure

FRFV Method – Calculations
Kadd = d * Fy-z or
Kadd = KJCz * Fy + (COPy – KJCy) * Fz

50

Massless Segmental Model
To obtain the knee adduction moment, calculations had to be made to obtain the reaction
forces and moments at the ankle.

These reaction forces and moments were then used to

determine the knee adduction moment.

This is a statically determinant model, and the

calculation of the knee adduction moment does not require knowledge of any value before or
after the frame of calculation.

Massless Segmental Model – Terms
Kadd – Knee Adduction Moment
Fy – Force component in the Y – direction (medial/lateral)
Fz – Force component in the Z – direction (vertical)
AJCy – Y component of the location of the ankle joint center
AJCz – Z component of the location of the ankle joint center
KJCy – Y component of the location of the knee joint center
KJCz – Z component of the location of the knee joint center
COPy – Y component of the center of pressure
ARFy – Y component of the ankle reaction force
ARFz – Z component of the ankle reaction force
Aabd – Ankle abduction moment
KRFy – Y component of the knee reaction force
KRFz – Z component of the knee reaction force

51

Massless Segmental Model – Ankle Force and Moment Calculations

∑ FAJC = 0
Fy + ARFy = 0
ARFy = − Fy
Fz + ARFz = 0
ARFz = − Fz

∑ M AJC = 0
Aabd + AJC z * Fy + ( AJC y − COPy ) * Fz = 0
Aabd = − AJC z * Fy − ( AJC y − COPy ) * Fz
Massless Segmental Model – Knee Force and Moment Calculations

∑ FKJC

=0

ARF y + KRF y = 0
KRF y = − ARF y
ARF z + KRF z = 0
KRF z = − ARF z

∑ M KJC

=0

Aabd + ( KJC z − AJC z ) * ARF y + ( KJC y − AJC y ) * ARF z = 0
K add = − ( KJC z − AJC z ) * ARF y + ( KJC y − AJC y ) * ARF z

52

Quasi-Static Segmental Model
The addition of segment masses and their locations as defined by Dempster (1955) is
what differentiates the quasi-static segmental model.

This model still does not taken into

account acceleration of the limbs, but it does account for how the mass of each segment with
respect to gravity acts on the system.

Quasi-Static Segmental Model -- Terms

COPy – Y Component of the Center of Pressure
Fy – Force component in the Y-direction
Fz – Force component in the Z-direction
g – Gravitational Acceleration
AJCy – Y Component of the Ankle Joint Center
AJCz – Z Component of the Ankle Joint Center
COMf,y—Y Component of the Foot Center of Mass
COMf,z—Z Component of the Foot Center of Mass
Mf—Mass of the Foot Segment
KJCy – Y Component of the Location of the Knee Joint Center
KJCz – Z Component of the Location of the Knee Joint Center
COMs,y—Y Component of the Shank Center of Mass
COMs,z—Z Component of the Shank Center of Mass
Ms—Mass of the Shank Segment
ARFy – Y Component of the Ankle Reaction Force
ARFz – Z Component of the Ankle Reaction Force

53

Aabd – Ankle Abduction Moment
KRFy – Y Component of the Knee Reaction Force
KRFz – Z Component of the Knee Reaction Force
Kadd – Knee Adduction Moment
Quasi-Static Segmental Model – Ankle Force and Moment Calculations

∑ FAJC = 0
Fy + ARFy = 0
ARFy = − Fy
Fz + M f * g + ARFz = 0
ARFz = − Fz − M f * g

∑ M AJC = 0
Aabd + ( AJC y − COPy ) * Fz + AJC z * Fy
+ ( AJC y − COM f , y ) * ( M f * g ) = 0
Aabd = −( AJC y − COPy ) * Fz − AJC z * Fy
− ( AJC y − COM f , y ) * ( M f * g )

54

Quasi-Static Segmental Model – Knee Force and Moment Calculations

∑ FKJC = 0
ARFy + KRFy = 0
KRFy = − ARFy
ARFz + KRFz + M s * g = 0
KRFz = − ARFz − M s * g

∑ M KJC = 0

K add − Aabd + ( KJC z − AJC z ) * ARFy + ( KJC y − AJC y ) * ARFz
+ ( KJC y − COM s, y ) * ( M s * g ) = 0
K add = Aabd − ( KJC z − AJC z ) * ARFy − ( KJC y − AJC y ) * ARFz
− ( KJC y − COM s, y ) * ( M s * g )
Full Dynamic Segmental Model
This model is the most accurate because it takes into account everything that the massless
and quasi-static models do, but it adds the effects of linear and angular acceleration of the foot
and shank segments to the system. This is the only model that requires knowledge of segment
motion before and after the frame being calculated in order to compute the adduction moment.
This is because the model takes a running average of the acceleration between the frames before
and after the frame in question.
Full Dynamic Segmental Model – Terms
COPy – Y Component of the Center of Pressure
Fy – Force component in the Y-direction
Fz – Force component in the Z-direction
g – Gravitational Acceleration
AJCy – Y Component of the Ankle Joint Center
55

AJCz – Z Component of the Ankle Joint Center
COMf,y —Y Component of the Foot Center of Mass
COMf,z —Z Component of the Foot Center of Mass
Mf —Mass of the Foot Segment
LAf,y – Linear Acceleration of the foot segment in the Y - Direction
LAf,z – Linear Acceleration of the foot segment in the Z - Direction
If— Foot Segment Moment of Inertia

α f,x—Angular Acceleration of the Foot about the X-Axis
KJCy – Y Component of the Location of the Knee Joint Center
KJCz – Z Component of the Location of the Knee Joint Center
COMs,y—Y Component of the Shank Center of Mass
COMs,z—Z Component of the Shank Center of Mass
Ms—Mass of the Shank Segment
LAs,y – Linear Acceleration of the foot segment in the Y - Direction
LAs,z – Linear Acceleration of the foot segment in the Z - Direction
If—Shank Segment Moment of Inertia

α s,x—Angular Acceleration of the Shank about the X-Axis
ARFy – Y Component of the Ankle Reaction Force
ARFz – Z Component of the Ankle Reaction Force
Aabd – Ankle Abduction Moment
KRFy – Y Component of the Knee Reaction Force
KRFz – Z Component of the Knee Reaction Force
Kadd – Knee Adduction Moment

56

Full Dynamic Segmental Model – Ankle Force and Moment Calculations

∑ FAJC = M f * LA f
Fy + ARFy = M f * LA f , y
ARFy = − Fy + Mf * LAf , y
Fz − M f * g + ARFz = M f * LA f , z
ARFz = M f * LA f , z − Fz − M f * g

∑ M COMf

= I f *α f , x
Aabd + (COM f , y − COPy ) * Fz + COM f , z * Fy
+ ( AJC y − COM f , y ) * ARFz + ( AJC z − COM f , z ) * ARFy = I f * α f , x
Aabd = I f * α f , x − (COM f , y − COPy ) * Fz − COM f , z * Fy −

( AJC y − COM f , y ) * ARFz − ( AJC z − COM f , z ) * ARFy

57

Full Dynamic Segmental Model – Knee Force and Moment Calculations

∑ FKJC = M s * LAs
ARF y + KRFy = M s * LAs, y
KRFy = M s * LAs, y − ARF y
ARFz + KRFz + M s * g = M s * LAs, z
KRFz = M s * LAs, z − ARFz − M s * g

∑ M COM

s

= I s * α s, x

K add − Aabd + (COM s, z − AJC z ) * ARF y + (COM s, y − AJC y ) * ARFz +
( KJC z − COM s, z ) * KRFy + ( KJC y − COM s, y ) * KRFz = I s * α s, x
K add = I s * α s, x + Aabd − (COM s, z − AJC z ) * ARF y
− (COM s, y − AJC y ) * ARFz − ( KJC z − COM s, z ) * KRFy
− ( KJC y − COM s, y ) * KRFz
Comparison of Segmental Models from One Trial of One
Subject
0.5
0.4
Moment (Nm/kg)

0.3
0.2

Dynamic
Quasi Static
Massless
FRFV

0.1
0
-0.1
-0.2
-0.3
-0.4 0

0.2

0.4

0.6

Time (s)
Figure 3.6 – Comparison of segmental models from one trial of one subject
58

In order to calculate the knee adduction moment from the dynamic segmental model, a lower
body marker had to be developed to estimate joint centers.
Marker Set
Ten markers were placed on body landmarks for static measurements and 8 retroreflective markers were used for the dynamic measurements. Table 3.1 lists the markers and
whether they are used for static only or static and dynamic models. The marker set chosen here
was similar to the set used by Fisher et. al. and Mundermann et. al. to determine lower body joint
segments for use in an inverse dynamics model (Figure 3.6) (Fisher et al., 2007; Mundermann,
Dyrby, Hurwitz, Sharma, & Andriacchi, 2004).
Table 3.1 –Marker Locations and Names for Static and Dynamic Trials
Marker Name
Right Superior Iliac Spine
Right Greater Trochanter
Right Posterior Thigh
Right Lateral Knee Joint
Right Medial Knee Joint
Right Posterior Calf
Right Lateral Malleolus
Right Medial Malleolus
Right Lateral Aspect of the Calcaneus
Right 5th Metatarsal Head

Abbreviation
RSIS
RGTR
RPOT
RLKJ
RMKJ
RPOC
RLMA
RMMA
RLAC
R5MH

59

Static
X
X
X
X
X
X
X
X
X
X

Dynamic
X
X
X
X
X
X
X
X

Side View

Anterior View

Posterior View

RSIS

RSIS
RGTR

RPOT

RPOT
RLKJ

RLKJ

RMKJ

RMKJ
RPOC

RPOC

RLMA

R5MH

RMM

RLMA

RLAC
RMMA

R5MH

Figure 3.7 – Anterior, Lateral, and Posterior Views
of the Anatomical Model with Marker Set

60

RLAC

Physical Markers
This section describes the marker set used to generate segments for the inverse dynamics model.
Figure 3.7 illustrates the anatomical skeletal landmarks that the marker set uses for reference.

RSIS

RGTR

RLKJ

RMKJ

RLMA
RMMA
RLAC
R5MH

Figure 3.8 – Anatomical skeletal landmark references
for the marker set

61

Superior Iliac Spine (RSIS) – This anatomical location is located at the top of the pelvic iliac
crest. This bony landmark is found by palpation after finding the anterior superior iliac spine on
the pelvis. From this landmark, the superior iliac spine is found by moving up the pelvic girdle
in a posterior fashion on the border of the iliac crest until the top is reached. The marker
attached to this landmark was not used in calculations, but was used as a check for correct
direction of motion.

Greater Trochanter (RGTR) – In the anatomical position, the greater trochanter is the
protrusion on the femur that is lateral to the femoral head. To locate the landmark, the subject
would internally and externally rotate their leg with their foot off the ground while the
investigator placed their hand on the hip and felt for a bony prominence. The marker was then
attached to this location.

Lateral and Medial Knee Joint Line (RLKJ, RMKJ) – To find the lateral knee joint line, the
subject sat in a chair and oriented themselves such that their thighs would be parallel to the floor
and the knee joint flexed at a 90 degree angle. A small mark was placed where the lateral aspect
of the knee rotation line appeared to be. To confirm if the mark was on the correct location, the
subject extended and flexed their leg about their knee joint without moving the thigh. If the
marker moved in an arc during this visual inspection, it was not at the knee joint line. The
marker was then adjusted until it only rotated and did not translate when the leg was flexed and
extended. This method was repeated for the medial aspect of the knee. These two markers
defined the rotation line of the knee. The midpoint between these two markers was important as
it was identified as the knee joint center.

62

Posterior Thigh (RPOT) – This location was not an anatomical landmark, but the marker that
was placed there allowed for a better visual understanding of how the gait cycle was progressing
and the direction of movement. The marker position was located by finding the midpoint
between the greater trochanter marker and the lateral knee joint line marker while the subject
was standing with their feet together. When the midpoint was found, an imaginary line was
traced around to the posterior side of the thigh. The marker was then placed at that position.

Medial Malleolus (RMMA) – The medial malleolus was round by palpating the inner portion of
the ankle joint. This landmark was used to help determine the width of the ankle joint. This
marker was not used during dynamic testing because it became occluded from the cameras
during the swing phase of the opposite leg.

Lateral Malleolus (RLMA) – The lateral malleolus was located in the same manner as the
medial malleolus. It is the bony landmark that protruded the furthest away from the ankle on the
lateral inferior border.

This mark was used in conjunction with the medial malleolus to

determine the location of the center of the ankle joint during the static trial. The ankle joint
center was the midpoint between the medial and lateral malleoli. The ankle joint center was
calculated from the lateral malleolus marker during dynamic trials.

Posterior Calf (RPOC) – The posterior calf marker was placed in the same manner that the
posterior thigh marker was placed. The midpoint between the lateral knee joint line and lateral
malleolus markers was determined and a marker was placed on the posterior portion of the lower

63

leg. Again, this marker was only used to visualize the leg and was not used in subsequent
calculations.

Lateral Aspect of the Calcaneus (RLAC) – This marker was applied directly to the footwear.
The marker was applied to the lateral heel of the shoe, as the calcaneus is typically covered by a
shoe. This marker was placed lateral to the approximate location of the center of the calcaneus
when visualized from the lateral side of the body. This marker placement helped define the foot
segment in subsequent analyses.

th

5

Metatarsal Head (R5MH) – This marker was applied to the outside of the shoe because

metatarsal heads were covered by the footwear. This landmark was found by palpating the 5

th

metatarsal until the distal portion of the digit was found. This marker also helped define the foot
nd

segment by being the reference marker for the 2

metatarsal head.

Virtual Markers
These next marker descriptions were virtual markers. This means that the actual position of the
marker was relative to locations of physical markers and had to be referenced to a physical
marker.
Hip Joint Center – In the current study, the hip joint was idealized as being located horizontally
and collinear with the marker placed on the greater trochanter. This marker was then offset by a
value that was determined by estimating the distance from the greater trochanter to the
acetabulum by using a ruler.

64

Knee Joint Center – The knee joint center was also a virtual marker. This marker was idealized
as being horizontal and collinear with the lateral knee joint line. It was offset by half of the
length between the lateral and medial knee joint line markers. The amount of offset was
determined by measuring the distance between the lateral and medial knee joint line markers
during the static trial and was verified using an anthropometer

Ankle Joint Center – The ankle joint center was also a virtual marker. This marker was
estimated to be horizontal and collinear with the lateral malleolus marker. As with the knee joint
center, this offset dimension was determined by measuring the distance between the lateral and
medial malleolus markers during the static trial and was verified using an anthropometer.

2

nd

Metatarsal Head – The position of the second metatarsal head was estimated by using the

position data from the fifth metatarsal head and the ankle joint center. The anterior/posterior and
vertical components of the fifth metatarsal head were used, while the medial/lateral value of the
ankle joint center was used to estimate this marker.
Materials and Methods
A Vicon Motion Analysis System (Vicon Inc., Oxford, UK) was used to capture
positional data at 100Hz from a skin mounted marker set. Reference videos were captured at
50Hz using a Basler 602f digital video camera. Force data were collected at 1000Hz using an
Advanced Mechanical Technology Incorporated AMTI force platform Model OR6-7-1000
(Advanced Mechanical Technology Incorporated, Watertown, MA). All data were recorded and
time-synchronized to a personal computer through the Vicon Nexus program. Force plate data
were first sent through an AMTI SGA6-4 Signal Conditioner/ Amplifier, to the A2D converter

65

board, to the Vicon MX Control Unit, through the Vicon MX Ultranet unit. The force platform
collected six channels of data: forces Fx, Fy, & Fz, and moments Mx, My, & Mz. The three
motion cameras fed directly into the MX Ultranet unit and into the computer. The motion
cameras each had a resolution of 0.3 Megapixels. Each camera recorded a two-dimensional
grayscale image of the retro-reflective markers which were placed on the body. Calibration of
the system allowed for recreation of the three-dimensional marker motion from these twodimensional grayscale images. Because the lab has only one force plate embedded in the raised
track, markers were only placed on the right side of the body. The right leg was also selfidentified as the dominant leg for each subject.
Test Subject Descriptions
Eight healthy subjects (4 male, 4 female) with no history of lower body surgery or other
recent lower body problems and the needed shoe sizes (men’s size 9 and women’s size 8) were
recruited for this study. The height and mass of each subject can be found in Table 3.2
Table 3.2 – Subject Demographics
Subject Demographics
Subject Code
Height (m) Weight (kg)
Subject F1
1.65
83.9
Subject F2
1.70
54.9
Subject F3
1.65
61.2
Subject F4
1.65
65.8
Subject M1
1.78
86.2
Subject M2
1.83
59.0
Subject M1
1.78
72.3
Subject M2
1.88
68.0

Each subject was tested in seven pairs of shoes designed by Merrell (Wolverine World Wide,
Rockford, MI) specifically for this battery of tests. The shoe design was based off the Merrell
66

Full Pursuit model trail running shoe. The stiffness values of the shoes were measured using an
Asker C Durometer scale to measure the stiffness of the ethylene-vinyl acetate (EVA) midsole
material. The control shoe had a density of 53 on the Asker C durometer scale across the entire
midsole. The lateral stiffened shoe was divided into right and left halves geometrically. The
medial half of the midsole was a C 43 and the lateral half was a C 73 (Figure 3.9). The other five
prototype production quality shoes were called T-Form shoes. Each T-Form shoe had a stiff
medial component that extended from the heel to the medial arch of C 73 durometer. The
difference between these shoes was the changing location of a pocket of lateral stiffness at C 73
stiffness. T-Form Shoe 1 had the stiff lateral pocket in the posterior part of the lateral heel. The
pocket of stiffness moved anterior with each subsequent shoe, ending with T-Form Shoe 5
having the pocket in the most anterior position of any shoe (Figure 3.10).

67

Neutral Control Shoe

Full Lateral Stiffened Shoe

53

43

73

Figure 3.9 – Photo of Merrell Full Pursuit shoe style
for neutral control shoe and full lateral stiffened shoe
(top) and the location and Asker C Stiffness values
for the midsole configurations (bottom)

68

T-Form
Shoe 1

T-Form
Shoe 2

T-Form
Shoe 3

T-Form
Shoe 4

T-Form
Shoe 5

43
43

43

73

73

73

43
43
73

73

73

Figure 3.10 – Photo of the Merrell Full Pursuit shoe
style for the five T-Form shoes (top) and the location
and Asker C Stiffness values for the midsole
configurations (bottom)

69

Testing Procedures
Prior to any walking trials being performed, the scope of the study was explained to the
test subject. Informed consent was obtained from each subject. The next step was to apply
retro-reflective markers to the skin using 3M double-sided tape reinforced with athletic training
paper tape. After application of the markers, the subject was asked to stand straight with their
feet shoulder width apart and parallel to each other facing down the length of the track for a
static trial. This trial was performed to obtain widths of the knee and ankle joints. The subject
was then allowed to practice walking down the track at a self-selected pace so that they could hit
the force plate with their right foot. The subject was told not to look down at the force plate
during testing to ensure that they did not shorten or lengthen their natural stride to hit the plate.
Each subject performed at least three good walking trials in each shoe model (Figure 3.11).

Figure 3.11 – Subject walking in a TForm model shoe across the force plate

70

Data Analysis
After data collection was complete, each trial was processed through the Vicon Nexus
program. The three-dimensional positional marker data were combined with the force plate data
and stored in a .c3d file. The .c3d file was then opened in the Vicon BodyBuilder program. The
marker trajectories were then labeled and the dynamic model that was written by the author was
applied to the data. The knee adduction moment data could then be exported with units of
Newton-millimeters from the BodyBuilder program as a .txt file. These values were then
converted to Newton-meters and normalized by dividing this value by the mass of the subject.
The units of the normalized data were Newton-meters per kilogram body mass. The data were
then plotted (Figure 3.12) and the maximum peak knee adduction moments were recorded for
each trial.

Knee Adduction Moment vs. Time -- Subject F3 T-Form Shoe 1
Trial 3
0.50
Peak Knee Adduction Moment

Moment (Nm/kg)

0.40
0.30
0.20
0.10
0.00
0

0.1

0.2

0.3

0.4

0.5

-0.10
Time (s)
Figure 3.12 – A typical knee adduction moment vs. time plot

71

0.6

0.7

Results
As can be seen in Table 3.3, the range of average knee adduction moments was from 0.17
Nm/kg in the Neutral Shoe for Subject M3 to 0.81 Nm/kg in Subject M1 for T-Form Shoe 3.
Because there were no statistically significant differences (two-way ANOVA, α =.05, all p>.05)
due to the large standard deviation between subjects, trends were said to be changes in the knee
adduction moment greater than or equal to 0.05 Nm/kg.

These differences are similar to

differences found in comparable literature on knee adduction moment changes (Crenshaw et al.,
2000; Fisher et al., 2007; Kakihana, Torii et al., 2005; Kerrigan et al., 2002).
Table 3.3 – Average and Standard Deviation of the Knee Adduction Moment (Nm/kg) for Each
Subject
Subject F1
Footwear
Condition
Neutral
Lateral Stiff
T-Form 1
T-Form 2
T-Form 3
T-Form 4
T-Form 5

Average
0.29
0.31
0.35
0.32
0.32
0.32
0.30

Neutral
Lateral Stiff
T-Form 1
T-Form 2
T-Form 3
T-Form 4
T-Form 5

Average
0.75
0.57
0.73
0.76
0.81
0.75
0.80

Subject F3

Subject F4

S.D. Average S.D. Average S.D. Average S.D.
0.01
0.35
0.06
0.45
0.03
0.36
0.01
0.01
0.46
0.04
0.47
0.06
0.38
0.01
0.01
0.38
0.03
0.48
0.04
0.40
0.06
0.01
0.35
0.02
0.45
0.04
0.44
0.03
0.01
0.29
0.03
0.44
0.04
0.48
0.04
0.04
0.38
0.06
0.44
0.03
0.39
0.03
0.03
0.25
0.09
0.46
0.03
0.44
0.02

Subject M1
Footwear
Condition

Subject F2

Subject M2

Subject M3

Subject M4

S.D. Average S.D. Average S.D. Average S.D.
0.04
0.37
0.08
0.17
0.03
0.38
0.07
0.08
0.28
0.00
0.24
0.02
0.31
0.03
0.03
0.27
0.08
0.25
0.01
0.37
0.05
0.05
0.21
0.07
0.20
0.02
0.39
0.03
0.06
0.22
0.06
0.24
0.02
0.35
0.03
0.07
0.28
0.01
0.29
0.01
0.35
0.03
0.02
0.25
0.11
0.27
0.03
0.35
0.02

72

Neutral (Control) vs. Lateral Stiffened Shoe
Three of the eight subjects (F1, F3, and F4) had differences less than 0.05 Nm/kg
between the neutral and lateral stiffened shoes. In addition, two subjects (F2 and M3) had
increases in the adduction moment greater than 0.05 Nm/kg from the neutral shoe to the lateral
stiffened shoe. The remaining three subjects (M1, M2, and M4) had decreases in the adduction
moment greater than 0.05 Nm/kg.
Neutral (Control) vs. T-Form Shoe 1
Five of the eight subjects (F2, F3, F4, M1, and M4) had less than a 0.05 Nm/kg
difference between the control and T-Form Shoe 1. Two subjects (F1 and M3) had an increase
in the moment from the neutral shoe to T-Form Shoe 1, and one subject (M2) had a decrease in
the knee adduction moment.
Neutral (Control) vs. T-Form Shoe 2
Six of eight subjects (F1, F2, F3, M1, M2, and M4) had a knee adduction moment
difference of less than 0.05 Nm/kg between the neutral and T-Form Shoe 2. Subject F4 had an
increase in the moment greater than 0.05 Nm/kg from the neutral to T-Form Shoe 2, while
Subject M3 had a decrease in knee adduction moment between the neutral and T-Form Shoe 2.
Neutral (Control) vs. T-Form Shoe 3
Four of eight subjects (F1, F3, M1, and M4) did not have a change of greater than 0.05
Nm/kg between the two shoes. Two subjects (F4 and M3) had increases in the adduction
moment from the neutral shoe to T-Form Shoe 3 while two subjects (F2 and M2) had a decrease
in the adduction moment.

73

Neutral (Control) vs. T-Form Shoe 4
Six of the eight subjects (F1, F2, F3, F4, M1, and M4) had changes of less than 0.05
Nm/kg between the neutral and T-Form Shoe 4. Subject M2 had a decrease in the adduction
moment greater than 0.05 Nm/kg between the neutral shoe and T-Form Shoe 4, while Subject
M3 had an increase in the adduction moment.
Neutral (Control) vs. T-Form Shoe 5
Three subjects (F1, F3, and M4) had changes in the adduction moment less than 0.05
Nm/kg between the neutral shoe and T-Form Shoe 5. In addition, Subjects F4, M1, and M3 had
an increase in the adduction moment from the neutral shoe to T-Form Shoe 5. Two subjects (F2
and M2) showed a decrease in the adduction moment between the neutral shoe and T-Form Shoe
5.

Average Overall Knee Adduction Moment
As can be seen from Table 3.4, there was almost no difference in the overall average of
the knee adduction moment between any of the footwear conditions. The overall variation in
values was 0.02 Nm/kg.
Table 3.4 –Average Overall Knee Adduction Moment

Neutral
Lateral Stiff
T-Form 1
T-Form 2
T-Form 3
T-Form 4
T-Form 5

Knee Joint Adduction Moment (Nm/kg)
Average (+ Standard Deviation)
0.39 (+ 0.17)
0.38 (+ 0.11)
0.40 (+ 0.15)
0.39 (+ 0.18)
0.39 (+ 0.19)
0.40 (+ 0.15)
0.39 (+ 0.18)

74

Gendered Average Knee Adduction Moment
It can be seen from Table 3.5 that there was an increase of 0.05 Nm/kg in the average
female adduction moment from the neutral shoe to the lateral stiff shoe, but a decrease in the
male average knee adduction moment of 0.07 Nm/kg. There were no changes greater or equal to
0.05 Nm/kg between the neutral and any of the T-Form style shoes.
Table 3.5 – Gendered Average of the Knee Adduction Moment
Female Knee Joint Adduction Moment (Nm/kg)
Average (+ Standard Deviation)
Neutral
0.36 (+ 0.07)
Lateral Stiff
0.41 (+ 0.08)
T-Form 1
0.40 (+ 0.06)
T-Form 2
0.39 (+ 0.06)
T-Form 3
0.38 (+ 0.09)
T-Form 4
0.38 (+ 0.05)
T-Form 5
0.36 (+ 0.10)

Neutral
Lateral Stiff
T-Form 1
T-Form 2
T-Form 3
T-Form 4
T-Form 5

Male Knee Joint Adduction Moment (Nm/kg)
Average (+ Standard Deviation)
0.42 (+ 0.18)
0.35 (+ 0.12)
0.41 (+ 0.19)
0.39 (+ 0.23)
0.41 (+ 0.24)
0.42 (+ 0.20)
0.42 (+ 0.23)

75

Discussion
The first purpose of this chapter was to develop a dynamic biomechanical model to
determine the knee joint adduction moment.

This was accomplished with a full dynamic

segmental model that accounted for the weight, orientation, and moment of inertia of each lower
limb segment. The adduction moment calculation was made with knowledge of the joint center
locations, segment orientation, the motion of the limbs, and the ground reaction force. This
model would be improved with the addition of markers to each body segment. This would allow
for more accurate representation of each segment and allow for additional joints to be
represented. Specifically, additional markers on the foot segment would help to better identify
the three main planes of subtalar motion to improve the foot kinematics. Also, additional
markers in the shank and leg segments could help to extend the model to measure the kinematics
and kinetics of the hip joint.
The second and third purposes of the study were to determine the effects of change in
midsole density patterns in production model shoes on the knee adduction moment. It was
hypothesized that the full lateral stiffened shoes would decrease the knee adduction moment as
compared to the neutral shoe condition in a manner similar to the changes documented
previously by others (Fisher et al., 2007). In addition, it was unclear whether the T-Form model
shoes would significantly change the knee adduction moment because of the shifting lateral
stiffened regions and the stiffened medial heel component. It was found that while some subject
specific responses to different shoe conditions were found, overall the knee adduction moment
was not significantly less in either the full lateral stiffened shoe or any T-Form shoe model, as
compared to the neutral control condition. This result was in contrast with the study done by
Fisher et. al. study which found differences between a neutral control shoe and a lateral stiffened

76

experimental shoe (Fisher et al., 2007). The effects of the lateral stiffened shoe may have been
mitigated by an outsole configuration in the model used in this study versus standing directly on
the midsole in the 2007 Fisher et. al. study. In addition, when changes in the knee adduction
moment were grouped by gender, there was an increase in the moment in the lateral stiffened
shoe as compared to the neutral shoe for female subjects of 0.05 Nm/kg, but an average decrease
of 0.07 Nm/kg between the neutral shoe and lateral stiffened shoe for male subjects. These
results were not expected, as gender-specific differences in the knee adduction moment in similar
conditions were not noted in any prior studies (Kerrigan, Riley, Nieto, & Della Croce, 2000) and
the fact that subjects from both genders were used in the studies documenting decreases in
adduction moments using lateral wedges and prototype lateral stiffened shoes (Crenshaw et al.,
2000; Fisher et al., 2007; Kakihana et al., 2004; Kerrigan et al., 2002).
It is possible that lack of difference in the knee adduction moment between the neutral
shoe and any of the T-Form shoes could be attributed to the stiffened medial heel for each
design. It has been shown that a medial wedge can effectively increase the adduction moment,
decreasing the lateral compartment knee load and mitigating pain in people with lateral
compartment knee osteoarthritis (Gross & Hillstrom, 2008). It is possible that stiffening the
medial heel of the T-Form shoes helped cancel out the effect of the various laterally stiffened
portions of the shoes.
Even though no overall change was found between the neutral and lateral stiffened
production shoe condition in subjects without medial compartment knee osteoarthritis, the
concept of changing midsole stiffness to elicit a biomechanical response could still have utility.
Variable density midsole shoes are sometimes used to prevent or limit the amount of pronation in
the form of motion control footwear. If proper density limits and locations could be determined,

77

it could lead to more comfortable footwear interventions to biomechanical problems. It has been
said that one reason for a subject to not continuing to wear an orthotic is because of the
discomfort of the intervention method (Kerrigan et al., 2002). If variable density midsole shoes
were to be explored further, it may lead to more comfortable orthotic intervention methods.

78

REFERENCES

79

References
Ahlback, S. (1968). Osteoarthrosis of the knee. A radiographic investigation. Acta Radiol Diagn
(Stockh), Suppl 277:277-272.
Crenshaw, S. J., Pollo, F. E., & Calton, E. F. (2000). Effects of lateral-wedged insoles on kinetics
at the knee. Clin Orthop Relat Res(375), 185-192.
Davis, M. A. (1988). Epidemiology of osteoarthritis. Clin Geriatr Med, 4(2), 241-255.
Felson, D. T., Zhang, Y., Hannan, M. T., Naimark, A., Weissman, B., Aliabadi, P., et al. (1997).
Risk factors for incident radiographic knee osteoarthritis in the elderly: the Framingham
Study. Arthritis Rheum, 40(4), 728-733.
Fisher, D. S., Dyrby, C. O., Mundermann, A., Morag, E., & Andriacchi, T. P. (2007). In healthy
subjects without knee osteoarthritis, the peak knee adduction moment influences the
acute effect of shoe interventions designed to reduce medial compartment knee load. J
Orthop Res, 25(4), 540-546.
Gross, K. D., & Hillstrom, H. J. (2008). Noninvasive Devices Targeting the Mechanics of
Osteoarthritis. Rheumatic Disease Clinics of North America, 34, 755-776.
Kakihana, W., Akai, M., Nakazawa, K., Takashima, T., Naito, K., & Torii, S. (2005). Effects of
laterally wedged insoles on knee and subtalar joint moments. Arch Phys Med Rehabil,
86(7), 1465-1471.
Kakihana, W., Akai, M., Yamasaki, N., Takashima, T., & Nakazawa, K. (2004). Changes of
joint moments in the gait of normal subjects wearing laterally wedged insoles. Am J Phys
Med Rehabil, 83(4), 273-278.
Kakihana, W., Torii, S., Akai, M., Nakazawa, K., Fukano, M., & Naito, K. (2005). Effect of a
lateral wedge on joint moments during gait in subjects with recurrent ankle sprain. Am J
Phys Med Rehabil, 84(11), 858-864.
Kerrigan, D. C., Lelas, J. L., Goggins, J., Merriman, G. J., Kaplan, R. J., & Felson, D. T. (2002).
Effectiveness of a lateral-wedge insole on knee varus torque in patients with knee
osteoarthritis. Arch Phys Med Rehabil, 83(7), 889-893.
Kerrigan, D. C., Riley, P. O., Nieto, T. J., & Della Croce, U. (2000). Knee joint torques: a
comparison between women and men during barefoot walking. Arch Phys Med Rehabil,
81(9), 1162-1165.

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Kuroyanagi, Y., Nagura, T., Matsumoto, H., Otani, T., Suda, Y., Nakamura, T., et al. (2007).
The lateral wedged insole with subtalar strapping significantly reduces dynamic knee
load in the medial compartment gait analysis on patients with medial knee osteoarthritis.
Osteoarthritis Cartilage, 15(8), 932-936.
Mow, V. C., & Huiskes, R. (2005). Basic Orthopaedic Biomechanics and Mechano-Biology (3rd
ed.). Philadelphia: Lippincott Williams & Wilkins.
Mundermann, A., Dyrby, C. O., Hurwitz, D. E., Sharma, L., & Andriacchi, T. P. (2004).
Potential strategies to reduce medial compartment loading in patients with knee
osteoarthritis of varying severity: reduced walking speed. Arthritis Rheum, 50(4), 11721178.
Prodromos, C. C., Andriacchi, T. P., & Galante, J. O. (1985). A relationship between gait and
clinical changes following high tibial osteotomy. J Bone Joint Surg Am, 67(8), 11881194.
Wells, R. P. (1981). The projection of the ground reaction force as a predictor of internal joint
moments. Bull Prosthet Res, 10-35, 15-19.
Winter, D. A. (2005). Biomechanics and Motor Control of Human Movement (3rd ed.).
Hoboken: John Wiley & Sons, Inc.

81

CHAPTER 4—EVALUATION OF A NOVEL FOOTWEAR DESIGN AND ITS
EFFECTS ON REAR-FOOT MOTION AND MECHANICAL STIFFNESS
COMPARISON TO COMMERCIAL MOTION CONTROL SHOES
Introduction
Excessive motion of the foot during gait has been linked to several problems in the lower
body. Specifically, excessive pronation and supination have been suggested to be a possible
cause of patellofemoral pain syndrome (Duffey, Martin, Cannon, Craven, & Messier, 2000;
Tiberio, 1987), Achilles tendon issues (Clement & Taunton, 1981; McCrory et al., 1999; Smart,
Taunton, & Clement, 1980), plantar fasciitis (Martin et al., 2001), and shin splints (Viitasalo &
Kvist, 1983). In the literature, the terms pronation and supination are often used interchangeably
with eversion and inversion, respectively. It is often unclear whether authors are using these
terms to describe triplanar or single plane rotations of the foot or heel. In this study, pronation
and supination refer to rotations about the anteroposterior axis of the rear-foot, as has been the
case in previous studies (Perry & Lafortune, 1995; Stacoff, Kalin, & Stussi, 1991). During gait,
the initial motion of pronation and supination takes place in the heel or calcaneus region of the
foot. Supination of the foot is shown in Figure 4.1a and pronation of the foot is shown in Figure
4.1b.

b

a

Figure 4.1—Anterior view of supination (a) and pronation (b)

82

Orthotic interventions and footwear alterations have been used as methods to modify the
amount of pronation or supination. Stacoff et. al., in a review of orthotic insert studies, reported
that orthotics had positive effects (i.e. less pain, better motion control) at rates between 70% and
80% on injured runners (Stacoff et al., 2000).

When attempting to modify pronation or

supination of the foot, interventions in the posterior part of the shoe are frequently used. Medial
stiffening or posting is a typical example of an orthotic design used to treat excessive pronation.
Conversely, lateral stiffening or posting may be used to treat excessive supination. Current
literature provides two suggestions for the efficacy of orthotics. First, the orthotic may change
the anatomical orientation of the foot and ankle structure (Stacoff et al., 2000). Second, the
orthotic may stimulate subcutaneous receptors in the foot and cause a change in the neural
pathway, thereby altering the progression of the step (Feuerbach, Grabiner, Koh, & Weiker,
1994). In actuality, it may be a combination of these two theories.
Using knowledge of successful orthotic designs, the Bates division of Wolverine World
Wide, Inc. (Rockford, MI) developed a novel insole technology designed to mitigate excessive
pronation or supination. This design was called the Wave Disk Comfort Technology. The Wave
Disk Comfort Technology, or WDCT, was designed to be a user-operated mechanism for people
to improve comfort and correct excessive rear-foot motion (Figure 4.2). Top and bottom views
of a ring can be found in Figures 4.3 and 4.4. The bottom of the ring contains the disk insertion
directions (Figure 4.5). If a user desired to have a strictly anti-pronation setting, they would line
up the P arrow on the disk with the arrow on the insole. The disks fit snugly onto a center post in
the heel of the modified insole. The post was not designed to carry a load, as it was not in
contact with the ground when the heel area was loaded. The post provided an attachment area to

83

ensure the stability of the WDCT disk. This study represents the first biomechanical evaluation
of this product.

Figure 4.2 – Posterior view of the modified insert with a Wave Disk (left) and the control insert
(right)

Figure 4.3—Top view of a red Wave Disk

84

Figure 4.4—Bottom view of a blue Wave Disk

Figure 4.5—Bottom view of a right WDCT insole with Wave Disk placed in the anti-pronation
setting
The disks work on a simple support principle. The peaks of the insole correspond with
the valleys of the WDCT rings and vice versa. The 8 peaks on the WDCT ring are 5 different
heights, decreasing in height around half the circumference (Figure 4.6 and Figure 4.7). The ring
design allows a user to determine the orientation of the directionally labeled disk. For persons
who would be considered excessive pronators, the disks would be aligned in the insole as shown

85

in Figure 4.6. For a person that needs more heel support on the lateral side, the disks would be
placed in the position as shown in Figure 4.7. The additional support is believed to bias an
individual so that they favor an altered rear-foot position; and as a result, limit the amount of
excessive rear-foot motion.
The other two extremes that the WDCT rings could be placed in were designated Firm
(F), and Regular (R), and would place the tallest peak at the rear of the heel or the anterior
portion of the heel, respectively. It is noted that these two positions represented a comfort setting
rather than a biomechanically-active setting, and therefore were not tested for their rearfoot
motion-limiting capabilities. It should be noted that it would also be possible for the WDCT
rings to be placed in a manner such that the user would experience elements of anti-pronation
and firm, anti-pronation and regular, anti-supination and firm, and anti-supination and regular.
Biomechanical evaluations of these settings were not performed during this round of testing as
the settings were not thought to provide maximum anti-pronation or anti-supination
characteristics because of the offset orientation of the technology.

Figure 4.6—Posterior view of the red Wave Disks in the anti-pronation setting

Figure 4.7—Posterior view of the red Wave Disks in the anti-supination setting
The purpose of this study was to determine if the novel WDCT footwear technology
could alter the amount of rear-foot pronation or supination during gait. It was hypothesized that
the anti-pronation setting would reduce rear-foot pronation as measured in human subjects

86

during walking gait. In addition, it was hypothesized that the amount of medial stiffness, as
measured in the heel region of the insole, would exceed the stiffness in the lateral facet when the
disc was placed in the anti-pronation setting. The anti-supination setting was thought to exhibit
similar alterations in rear-foot motions and insole stiffness, but affecting the opposite side of the
foot and insole heel region. The results of this study may influence future orthotic and insert
design towards designs that are able to be changed by the end user for comfort and function,
rather than having just one permanent setting.

Materials and Methods

Footwear and Wave Disk Comfort Technology
Two different pairs of Bates Stock Number B06-21 boots were used for these tests
(Figure 4.8). Both pairs were men’s size 8 boots with medium width characteristics and the
same outer shell. The control pair of boots was unmodified, while the test pair of boots was
modified in order to be able to house the Wave Disk Comfort Technology rings with a special
insole that housed the rings. The modification was most noticeable in the heel region, as extra
vertical room was made for allowable clearances between the foot and the boot while factoring
in the extra height of the circular inserts.

87

Figure 4.8 – Bates B06-21 Boots

For this study, the disks were tested in their anti-pronation and anti-supination settings.

Subjects
Two male subjects without a history of foot or ankle surgery or other orthopaedic
problems were tested in three experimental conditions – a control setting, an anti-supination
setting, and an anti-pronation setting.

Kinematics and Kinetics
Ground reaction force data were collected at 1000 Hz using an AMTI OR 6-7 force plate
(Advanced Mechanical Technology, Inc., Watertown, MA) embedded in a raised track in the
Orthopaedic Biomechanics Gait Laboratory. Kinematic ankle data were collected using a three
camera Vicon motion capture system (Vicon, OMB plc., Oxford, UK) at 100 Hz. In order to
evaluate ankle motion, retro-reflective marker sets were attached posterior to the calcaneus

88

(Figure 4.9) and at the base of the third metatarsal on the dorsal (top) side of the foot (Figure
4.10). Each marker was 9 mm in diameter.

Figure 4.9—Marker set for the posterior calcaneus

Figure 4.10—Marker set for the dorsal forefoot
The attachment sites for the custom marker arrays were made by taping a flexible
receiving piece to the foot (Figure 4.11). This type of attachment is not suitable for measurement
of precise skeletal motion (Stacoff et al., 2000) because the arrays cannot be directly attached to
the bone, but is suitable to measure the motions of the posterior foot during the gait cycle. The
custom marker arrays were designed for ease of rear foot motion calculation. The Euler angle
measurement was based on defining the axis of rear-foot rotation that was defined to be

89

pronation and supination motions. The forward-placed marker set was important in discerning
the spacial orientation of the foot during the testing as well as defining the direction of motion
for visualization in the motion capture software.

Figure 4.11—Picture of the posterior calcaneus marker set placed in the flexible receiving piece

Kinetic and Kinematic Procedure
Subjects had their right foot taped such that the cup portion of the flexible receiving piece
was firmly in contact with both the skin posterior to the calcaneus and the base of the third
metatarsal (Figure 4.12a). Figure 4.12b shows how the marker sets appear before placing the
foot in the boot. Because of the size of the marker sets, the foot could not be placed inside of the
boot with the marker sets already attached. The subject then placed the boot over the taped foot
without the maker sets such that the insertion cups were visible through the holes that were cut in
the shoe. The marker sets were then placed in their subsequent locations and were checked to
ensure that the posts emanating from the foot did not contact the sides of the holes or the ground
in the shoes during walking gait (Figure 4.12c).

90

a

b

c
Figure 4.12—From top left, clockwise—(a) Right foot taped with
flexible receiving pieces exposed; (b) Right foot with marker sets
attached; (c) Right foot with boot and marker sets attached
Prior to performing recorded trials, the subject was allowed to practice walking at a selfselected pace along the raised track in order to ensure that the force plate would capture their
entire step without looking at the force plate. This was done to ensure that the subject did not
alter their gait in order to hit the force plate with either a short or long stride which would
produce data inconsistent with a normal walking gait pattern. Once the subject was comfortable
with the walking protocol, the rear-foot marker set was aligned such that the axis of rear-foot
motion was perpendicular to the plane of the marker set. This was done to easily measure the
rotation angle between the ground and rear-foot.

The kinetics and kinematics were then

collected using the force plate and motion capture system. Each subject performed at least three
walking trials in the control shoe, three in the test shoe with the disk in anti-pronation setting,
and three in the disk in the anti-supination setting. The control shoe was an unmodified version
91

of the Bates B06-21 boots. Analysis was performed using a custom program designed in the
Vicon BodyBuilder software package to measure the angle of the heel with respect to the
horizontal plane.
Changes in angular rotation were measured between different established events that
occur during a typical gait cycle. The measurements were made at three time intervals –
between maximum braking force and first peak vertical force, between first peak vertical force
and mid-stance, and between braking force and mid-stance. The maximum braking force occurs
in the anterior-posterior direction, while the first peak and mid-stance events are determined
from the vertical force component of the ground reaction force (Figure 4.13).

Force vs. Time
800

First Peak
Vertical Force

700
600

Force (N)

500

Anterior/
Posterior
Force
Vertical
Force

400
300

Midstance

200
100

Peak Braking Force

0
-100
-200
0

0.1

0.2

0.3

0.4

0.5

0.6

Time (s)
Figure 4.13 – Graphical Explanation of Standard Gait Events

92

Motions that occur after mid-stance were not considered because the majority of the force
that is being applied through the foot after mid-stance does not contact the area where the WDCT
is applied. The angles reported are the changes in the angle from the first event to the second
event (i.e. change in angle from braking to first peak). A positive value means a pronation
motion, while a negative value indicates a supination motion. Change in rear-foot angles were
measured until mid-stance because Hintermann and Nigg (1998) have documented that the foot
rotates from initial contact with the ground until the mid-stance phase of gait. Also, after midstance, the heel begins to rise off the ground, making the motion forward of the heel technology.

Mechanical Footwear Construction Testing Procedure
The impact of the WDCT on the physical properties of the insole was assessed via
mechanical testing. The right insole was taken out of the boot and tested in a custom stiffness
test using an Instron machine (Instron, Inc., Norwood, MA) and a rectangular indenter head
measuring 19mm x 42mm. This indenter head had a surface area that was exactly half of the
area used in a standard shoe impact test (ASTM F1976-06). The heel of the Wave Disk insole
was divided into two areas (medial and lateral) centered about the hole in the WDCT ring. The
two divisions are highlighted in the following figure (Figure 4.14).

93

Rectangular Indenter Head

Medial

Lateral

Figure 4.14 – Location of Instron stiffness testing using 19mm x 42mm rectangular indenter
head
Each part of the heel was loaded to 623 N and the amount of deformation of the
insole was measured. The stiffness was calculated by dividing the peak load by the peak
deformation. These tests were performed using the WDCT ring in anti-pronation and antisupination settings. For stiffness comparison purposes, the same stiffness test was performed on
other shoes that were considered to be motion control shoes based on their marketing and design.

Kinematic Results
As can be seen in all tables, the amount of inversion or eversion occurring was generally small in
magnitude. In Table 4.1, the amount of rotation from braking to the first peak vertical force was
described. Subject 1 exhibited more eversion in both the Anti-Pronation and Anti-Supination
setting when compared to the Control setting, but had less eversion in the Anti-Pronation setting

94

as compared to the Anti-Supination setting. Subject 2 did not exhibit a significant difference in
the average change in angle for any of the three test conditions.

Table 4.1 – Summary of angle change from braking to first peak
Subject 1
Shoe Condition
Control Trial 1
Control Trial 2
Control Trial 3
Red Disk, Anti-Pronation Trial 1
Red Disk, Anti-Pronation Trial 2
Red Disk, Anti-Pronation Trial 3
Red Disk, Anti-Supination Trial 1
Red Disk, Anti-Supination Trial 2
Red Disk, Anti-Supination Trial 3

Braking to First Peak
Change in Angle (deg.)
0.85
-2.48
0.67
2.22
0.00
1.05
2.60
0.44
2.56

Average
Change

Subject 2
Shoe Condition
Control Trial 1
Control Trial 2
Control Trial 3
Red Disk, Anti-Pronation Trial 1
Red Disk, Anti-Pronation Trial 2
Red Disk, Anti-Pronation Trial 3
Red Disk, Anti-Supination Trial 1
Red Disk, Anti-Supination Trial 2
Red Disk, Anti-Supination Trial 3

Braking to First Peak
Change in Angle (deg.)
-0.53
0.60
-0.75
0.15
-0.39
-0.24
-0.33
-0.19
-0.41

Average
Change

-0.32

1.09

1.87

-0.23

-0.16

-0.31

In the comparison of the angle change from the first peak to midstance phases of gait, Subject 1
had decreases in eversion in both the Anti-Pronation setting and the Anti-Supination setting
when compared to the Control setting. Subject 2 exhibited a small decrease in eversion in the

95

Anti-Pronation setting compared to the Control setting and no significant change in the
comparison between the Anti-Supination setting and the Control setting.

Table 4.2 – Summary of Angle Change from First Peak to Midstance
Subject 1
Shoe Condition
Control Trial 1
Control Trial 2
Control Trial 3
Red Disk, Anti-Pronation Trial 1
Red Disk, Anti-Pronation Trial 2
Red Disk, Anti-Pronation Trial 3
Red Disk, Anti-Supination Trial 1
Red Disk, Anti-Supination Trial 2
Red Disk, Anti-Supination Trial 3

First Peak to Midstance
Change in Angle (deg.)
-0.25
2.84
1.51
0.27
0.77
1.21
-3.86
-0.51
0.17

Average
Change

Subject 2
Shoe Condition
Control Trial 1
Control Trial 2
Control Trial 3
Red Disk, Anti-Pronation Trial 1
Red Disk, Anti-Pronation Trial 2
Red Disk, Anti-Pronation Trial 3
Red Disk, Anti-Supination Trial 1
Red Disk, Anti-Supination Trial 2
Red Disk, Anti-Supination Trial 3

First Peak to Midstance
Change in Angle (deg.)
0.57
1.93
1.23
2.26
1.07
-1.23
0.41
1.46
2.12

Average
Change

1.37

0.75

-1.40

1.24

0.70

1.33

Subject 1 had an increase in eversion when comparing the Anti-Pronation to the Control, and a
decrease when comparing the Anti-Supination to the Control. Subject 2 showed a decrease in
eversion from the Anti-Pronation setting as compared to the Control setting, and no change
between the Anti-Supination setting and the Control.

96

Table 4.3 – Summary of Angle Change from Braking to Midstance
Subject 1
Shoe Condition
Control Trial 1
Control Trial 2
Control Trial 3
Red Disk, Anti-Pronation Trial 1
Red Disk, Anti-Pronation Trial 2
Red Disk, Anti-Pronation Trial 3
Red Disk, Anti-Supination Trial 1
Red Disk, Anti-Supination Trial 2
Red Disk, Anti-Supination Trial 3

Braking to Midstance
Change in Angle (deg.)
0.60
0.36
2.18
2.49
0.77
2.26
-1.26
-0.07
2.73

Average
Change

Subject 2
Shoe Condition
Control Trial 1
Control Trial 2
Control Trial 3
Red Disk, Anti-Pronation Trial 1
Red Disk, Anti-Pronation Trial 2
Red Disk, Anti-Pronation Trial 3
Red Disk, Anti-Supination Trial 1
Red Disk, Anti-Supination Trial 2
Red Disk, Anti-Supination Trial 3

Braking to Midstance
Change in Angle (deg.)
0.04
2.53
0.48
2.41
0.68
-1.47
0.08
1.27
1.71

Average
Change

1.05

1.84

0.47

1.02

0.54

1.02

Mechanical Testing Results
As can be seen from Table 4.4, the insole with the Wave Disk performed as intended, with the
stiffness of the medial side in the anti-pronation setting being stiffer than the lateral side. The
technology is also properly working in the anti-supination setting, with the lateral side being
stiffer than the medial side.

97

Table 4.4 – Stiffness values for B06-21 Insole with Red Wave Disk
Disk
Setting

Description

Right B 06-21
AntiInsole, Red Wave
Pronation
Disk
Right B 06-21
AntiInsole, Red Wave
Pronation
Disk
Right B 06-21
AntiInsole, Red Wave
Supination
Disk
Right B 06-21
AntiInsole, Red Wave
Supination
Disk

Heel
Location

Input
Force
(N)

Displacement
(mm)

Medial

625.059

Stiffness
%
Values
Difference
(N/mm)

6.209

100.67
10.74%

Lateral

624.317

6.947

89.87

Medial

624.873

7.539

82.89
19.40%

Lateral

625.337

6.081

102.83

In order to understand how the stiffness of the anti-pronation setting values compared with
production model shoes designed to prevent excessive pronation, the same Instron test was
performed on three motion control shoes.
Table 4.5 – Stiffness Values for Select Motion Control Shoes
Shoe
Description

Heel
Location

Input
Force
(N)

Displacement
(mm)

Stiffness
Values
(N/mm)

Merrell Motion
Control

Medial

623.791

8.017

77.81

Merrell Motion
Control

Lateral

624.207

9.317

67.00

Saucony Grid
Stabil 6

Medial

624.115

9.003

69.33

Saucony Grid
Stabil 6

Lateral

623.466

10.252

60.82

Asics Gel
Foundation

Medial

623.744

8.092

77.08

Asics Gel
Foundation

Lateral

%
Difference

13.90%

12.28%

18.37%
623.791

9.914

98

62.92

The results of comparing the lateral and medial sides of the three motion control shoes were
similar, as a Merrell motion control shoe was 10.81 N/mm stiffer on the medial side as compared
to the lateral side, the Saucony Grid Stabil 6 was 8.51 N/mm stiffer medial, and the Asics Gel
Foundation VI was 14.16 N/mm stiffer medial.
Discussion
This study focused on how a novel footwear design affected pronation and supination
during walking and how the design mechanically compared to current rear-foot motion control
footwear. It was hypothesized that the footwear design would change the amount of rear-foot
motion compared to a control setting. Also, it was hypothesized that the new footwear design
would have similar medial and lateral stiffness characteristics as compared to commerciallyavailable motion control shoes. Motion control shoes have been shown to be effective at
reducing rear-foot motion (Cheung & Ng, 2007). It was found that the WDCT insole had similar
stiffness characteristics as commercial motion control shoes, but results from subject testing did
not indicate that the technology was effective at changing the amount of rear-foot motion during
walking. Even though there was a medial and lateral heel stiffness difference similar to motion
control shoes, both sides of the WDCT insole were stiffer than any motion control shoe.
Because the insole behaved as intended in the mechanical tests, it was thought that the
change in subtalar motion during the biomechanical tests may be attributed to the construction of
the boot. The boot upper, heel cup and medial mid-foot region of the boot were stiff, and did not
allow much rotational movement when the boot was laced. This feature may have limited the
effects of the WDCT in this particular style of boot. In this study, the change in angle was
measured with respect to the horizontal. If the study was to be repeated, it is suggested that the
angle of rear-foot motion should be measure with respect to the angle of the tibia instead of a

99

fixed horizontal. This would provide a consistent reference plane with respect to the rear-foot
motion and is similar to testing described in literature (Cheung & Ng, 2007; Stacoff et al., 2000).
Even though no difference was found in the rear-foot kinematics in this boot model, the
WDCT still may be a viable idea. If the technology was implemented in a mid-top or low top
shoe, the technology may be more effective because of the increase in subtalar mobility in
footwear with shorter upper construction. However, there is generally more vertical heel room in
a taller boot as compare to a low-top shoe, and without modifying the WDCT it could be difficult
to successfully put the disc in a lower profile type of footwear. Also, to fully validate the
technology, more subjects as well as subjects who have excessive pronation problems should be
tested. Further testing of the WDCT may reveal a decrease in excessive rear-foot motion during
the gait of certain individuals, possibly mitigating the risk of injury and discomfort.

100

REFERENCES

101

References
Cheung, R. T. H., & Ng, G. Y. F. (2007). Efficacy of motion control shoes for reducing
excessive rearfoot motion in fatigued runners. Physical Therapy in Sport, 8, 6.
Clement, D. B., & Taunton, J. E. (1981). A guide to the prevention of running injuries. Aust Fam
Physician, 10(3), 156-161, 163-154.
Duffey, M. J., Martin, D. F., Cannon, D. W., Craven, T., & Messier, S. P. (2000). Etiologic
factors associated with anterior knee pain in distance runners. Med Sci Sports Exerc,
32(11), 1825-1832.
Feuerbach, J. W., Grabiner, M. D., Koh, T. J., & Weiker, G. G. (1994). Effect of an ankle
orthosis and ankle ligament anesthesia on ankle joint proprioception. Am J Sports Med,
22(2), 223-229.
Martin, J. E., Hosch, J. C., Goforth, W. P., Murff, R. T., Lynch, D. M., & Odom, R. D. (2001).
Mechanical treatment of plantar fasciitis. A prospective study. J Am Podiatr Med Assoc,
91(2), 55-62.
McCrory, J. L., Martin, D. F., Lowery, R. B., Cannon, D. W., Curl, W. W., Read, H. M., Jr., et
al. (1999). Etiologic factors associated with Achilles tendinitis in runners. Med Sci Sports
Exerc, 31(10), 1374-1381.
Perry, S. D., & Lafortune, M. A. (1995). Influences of inversion/eversion of the foot upon impact
loading during locomotion. Clin Biomech (Bristol, Avon), 10(5), 253-257.
Smart, G. W., Taunton, J. E., & Clement, D. B. (1980). Achilles tendon disorders in runners--a
review. Med Sci Sports Exerc, 12(4), 231-243.
Stacoff, A., Kalin, X., & Stussi, E. (1991). The effects of shoes on the torsion and rearfoot
motion in running. Med Sci Sports Exerc, 23(4), 482-490.
Stacoff, A., Reinschmidt, C., Nigg, B. M., van den Bogert, A. J., Lundberg, A., Denoth, J., et al.
(2000). Effects of foot orthoses on skeletal motion during running. Clin Biomech (Bristol,
Avon), 15(1), 54-64.
Tiberio, D. (1987). The effect of excessive subtalar joint pronation on patellofemoral mechanics:
a theoretical model. J Orthop Sports Phys Ther, 9(4), 160-165.

102

CHAPTER 5 – SUMMARY AND FUTURE DIRECTIONS OF STUDY

This thesis documented methods for measuring biomechanical changes due to
modifications in footwear midsole characteristics. Shifts in the foot center of pressure related to
changes in midsole stiffness and varying locations of increased midsole stiffness were measured
using an insole plantar pressure measurement system. In addition, knee adduction moments and
their response to full lateral stiffened shoes, as well as five other experimental shoe designs were
measured by using a full dynamic biomechanical model. Also, a novel footwear technology
designed to alter rear-foot motion was evaluated using a motion capture system.
In Chapter 2, an insole pressure measurement system was used to measure shifts in the
center of pressure in response to variations in midsole stiffening characteristics. The two shoe
conditions that elicited changes in the center of pressure were the full medial stiffened and the
medial stiffened heel with a lateral stiffened forefoot. Both footwear conditions consistently
shifted the center of pressure toward the lateral border of the shoe. It was noted that when the
shift did occur, it happened in the midfoot region and not at the heel and toe-off areas. In
addition, even when the center of pressure shift did occur for a majority of steps, it did not occur
in each step. This result coincided with results documented in the literature (B. M. Nigg et al.,
2003). The shift in the center of pressure was not large, as large shifts in the center of pressure
may lead to ankle joint instability and possible injury (Fuller, 1999). The knowledge that
changes in midsole density can change the center of pressure could be important to provide an
alternative to uncomfortable orthotics.

If a patient is not comfortable with an orthotic

intervention, they are less likely to continue its use. If the same biomechanical end results from

103

an orthotic can be achieved through a less invasive method, such as a variable stiffness midsole,
patients may benefit by use of this more comfortable intervention.
In Chapter 3, a lower body biomechanical model was developed in order to measure knee
adduction moments. It is important to be able to manipulate this moment, as it has been shown
that decreasing the knee adduction moment can slow the onset or progression of medial
compartment knee osteoarthritis. People with knee osteoarthritis are 10 times for likely to have
osteoarthritis in the medial compartment compared to the lateral compartment (Ahlback, 1968).
Current methods for decreasing the adduction moment include lateral wedges and invasive
surgeries, such as the high tibial osteotomy. A study by Fisher et. al. documented that stiffening
the lateral half of the midsole by 20% and 50%, as compared to the medial half, decreased the
adduction moment in a fashion similar to a 4 degree lateral wedge (Fisher, Dyrby, Mundermann,
Morag, & Andriacchi, 2007). This result was important, but their footwear was experimental
and not a typical production quality shoe. The biomechanical model was used to evaluate the
knee adduction moment in eight healthy subjects (four male, four female) while walking over a
force plate in seven different shoe conditions with different midsole characteristics.

The

footwear conditions were a neutral control shoe with a constant density throughout, a full lateral
stiffened shoe, and five T-Form shoes with a stiff medial heel component and a changing
location of a pocket of lateral stiffness. All shoes were production quality. On average, no
significant change was found between the neutral control shoe and any of the six test shoes. The
overall variation in the average knee adduction moment for all subjects was 0.02 Nm/kg. One
interesting finding was that the adduction moment decreased on average from the neutral to the
lateral stiffened shoe in male subjects, but increased in female subjects. No differences were
found between the T-Form model shoes and the neutral control shoe. It was also interesting that

104

the neutral shoe generated a larger adduction moment in the male subjects than in the female
subjects. This was not expected, as gender-specific differences have not been noted in previous
literature. In fact, Kerrigan et. al. found no difference in the adduction moment between men
and women walking barefoot (Kerrigan, Riley, Nieto, & Della Croce, 2000). Even though no
overall change was found between the neutral control shoe and the other shoe conditions, the
concept of variable density midsole stiffness shoes could still have utility. Refining the midsole
design and stiffness characteristics could lead to effect subject specific biomechanical
interventions to not only medial compartment knee osteoarthritis, but lateral compartment
osteoarthritis as well.
In Chapter 4, a novel footwear technology was designed to alter rear-foot motion. The
design was compared mechanically to current rear-foot motion control footwear.

This

technology was called the Wave Disk Comfort Technology (WDCT). It was documented that
the WDCT insole had similar medial and lateral heel stiffness characteristics as commercial
motion control shoes, but subject testing did not indicate that it was effective in changing the
amount of rear-foot pronation in our subjects. Because the insole behaved as intended in
mechanical testing, it was surmised that the lack of significant results in the subject testing could
possibly be attributed to the commercial boot construction. The build of the boot may have been
such that there were rear-foot motion limiting characteristics built into the shoe via the stiffness
and rigidity of the heel cup and boot upper. Even though no significant differences were
documented in rear-foot motion in the current study on the WDCT, it may still have utility when
implemented in a low or mid top model boot. It may be more effective because there may be
more subtalar mobility in those types of footwear.

In fact, this design concept has been

implemented in a number of footwear models by Wolverine World Wide, Inc. The device is also

105

currently being re-designed to improve efficacy across a wider range of footwear designs. Based
on the studies performed for this thesis, we have suggested that it would be appropriate to test
more subjects, both male and female, and test subjects who might have excessive rear-foot
motion to fully evaluate this footwear technology. To date, only subjects considered “normal”
have been used in evaluations of such technologies.
This research was performed to develop biomechanical models and use them to begin to
evaluate changes in the midsole characteristics of footwear that is designed to elicit changes in
various gait patterns. The work often represented early studies in the Orthopaedic Biomechanics
Laboratories Gait Lab at MSU. The studies described here suggest changes in midsole stiffness
characteristics may be able to alter center of pressure and knee adduction moments in selected
test subjects.
In addition to developing baseline experimental methods to compare more completely the
effects of various footwear technologies in Wolverine World Wide, Inc. footwear versus its
competitors, this thesis has allowed this laboratory to begin research in the area of “whole–body”
biomechanics. In the process the Vicon Motion Capture system was purchased, the BodyBuilder
software and BodyLanguage modeling was learned, an insole pressure measurement has been
developed and used for gait evaluations, and future plans are to incorporate the measurement of
muscle activity through electromyography (EMG) data into these evaluation technologies. Such
information is of current interest in unstable shoe design to provide muscle toning (B. Nigg,
Hintzen, & Ferber, 2006) and barefoot running shoes to help strengthen foot and lower body
musculature (Lieberman et al., 2010).

106

REFERENCES

107

References
Ahlback, S. (1968). Osteoarthrosis of the knee. A radiographic investigation. Acta Radiol Diagn
(Stockh), Suppl 277:277-272.
Fisher, D. S., Dyrby, C. O., Mundermann, A., Morag, E., & Andriacchi, T. P. (2007). In healthy
subjects without knee osteoarthritis, the peak knee adduction moment influences the
acute effect of shoe interventions designed to reduce medial compartment knee load. J
Orthop Res, 25(4), 540-546.
Fuller, E. A. (1999). Center of pressure and its theoretical relationship to foot pathology. J Am
Podiatr Med Assoc, 89(6), 278-291.
Kerrigan, D. C., Riley, P. O., Nieto, T. J., & Della Croce, U. (2000). Knee joint torques: a
comparison between women and men during barefoot walking. Arch Phys Med Rehabil,
81(9), 1162-1165.
Lieberman, D. E., Venkadesan, M., Werbel, W. A., Daoud, A. I., D'Andrea, S., Davis, I. S., et al.
(2010). Foot strike patterns and collision forces in habitually barefoot versus shod
runners. Nature, 463(7280), 531-535.
Nigg, B., Hintzen, S., & Ferber, R. (2006). Effect of an unstable shoe construction on lower
extremity gait characteristics. Clin Biomech (Bristol, Avon), 21(1), 82-88.
Nigg, B. M., Stergiou, P., Cole, G., Stefanyshyn, D., Mundermann, A., & Humble, N. (2003).
Effect of shoe inserts on kinematics, center of pressure, and leg joint moments during
running. Med Sci Sports Exerc, 35(2), 314-319.

108

APPENDICES

109

APPENDIX A
DATA FROM CHAPTER 2

110

Red COP line
for the control
shoe

Blue COP line for
the comparison
shoe

A green box indicates
higher pressure in
comparison shoe than
control shoe. A larger
box means a greater
difference.

A red box indicates
lower
pressure
in
comparison shoe than
control shoe. A larger
box means a greater
difference.

Figure A.1 – Definitions of terms and visualizations for reading Parotec data

111

Figure A.2 – Subject M1 Data - Shoe 1 vs. Shoe 2

112

Figure A.3 – Subject M1 Data - Shoe 1 vs. Shoe 3

113

Figure A.4 – Subject M1 Data - Shoe 1 vs. Shoe 4

114

Figure A.5 – Subject M1 - Data Shoe 1 vs. Shoe 5

115

Figure A.6 – Subject M1 Data - Shoe 1 vs. Shoe 6

116

red) vs
(blue)
teps
bject

Figure A.7 – Subject M2 Data - Shoe 1 vs. Shoe 2

117

Figure A.8 – Subject M2 Data - Shoe 1 vs. Shoe 3

118

Figure A.9 – Subject M2 Data - Shoe 1 vs. Shoe 4

119

Figure A.10 – Subject M2 Data - Shoe 1 vs. Shoe 5

120

Figure A.11 – Subject M2 Data - Shoe 1 vs. Shoe 6

121

Figure A.12 – Subject M3 Data - Shoe 1 vs. Shoe 2

122

Figure A.13 – Subject M3 Data - Shoe 1 vs. Shoe 3

123

Figure A.14 – Subject M3 Data - Shoe 1 vs. Shoe 4

124

Figure A.15 – Subject M1 Data - Shoe 1 vs. Shoe 5

125

Figure A.16 – Subject M3 Data - Shoe 1 vs. Shoe 6

126

Figure A.17 – Subject F1 Data - Shoe 1 vs. Shoe 2

127

Figure A.18 – Subject F1 Data - Shoe 1 vs. Shoe 3

128

Figure A.19 – Subject F1 Data - Shoe 1 vs. Shoe 4

129

Figure A.20 – Subject F1 Data - Shoe 1 vs. Shoe 5

130

Figure A.21 – Subject F1 Data - Shoe 1 vs. Shoe 6

131

Figure A.22 – Subject F2 Data - Shoe 1 vs. Shoe 2

132

Figure A.23 – Subject F2 Data - Shoe 1 vs. Shoe 3

133

Figure A.24 – Subject F2 Data - Shoe 1 vs. Shoe 4

134

Figure A.25 – Subject F2 Data - Shoe 1 vs. Shoe 5

135

Figure A.26 – Subject F2 Data - Shoe 1 vs. Shoe 6

136

Figure A.27 – Subject F3 Data - Shoe 1 vs. Shoe 2

137

Figure A.29 – Subject F3 Data - Shoe 1 vs. Shoe 3

138

Figure A.29 – Subject F3 Data - Shoe 1 vs. Shoe 4

139

Figure A.30 – Subject F3 Data - Shoe 1 vs. Shoe 5

140

Figure A.31 – Subject F3 Data - Shoe 1 vs. Shoe 6

141

APPENDIX B
DATA FROM CHAPTER 3

142

Table B.1 – Knee Moment Data for Subject F1
Subject F1
Shoe
Trial #
Neutral
Neutral
Neutral
Lateral Stiff
Lateral Stiff
Lateral Stiff
T-Form 1
T-Form 1
T-Form 2
T-Form 2
T-Form 2
T-Form 3
T-Form 3
T-Form 3
T-Form 4
T-Form 4
T-Form 4
T-Form 5
T-Form 5
T-Form 5

1
2
3
1
2
3
1
2
1
2
3
1
2
3
1
2
3
1
2
3

Knee Adduction/Abduction Moment
Normalized Joint Moment (Nm/k g) Average Standard Deviation
0.29
0.29
0.01
0.28
0.29
0.32
0.01
0.31
0.30
0.31
0.35
0.01
0.35
0.34
0.33
0.01
0.32
0.32
0.32
0.32
0.01
0.32
0.33
0.31
0.31
0.37
0.04
0.32
0.29
0.28
0.30
0.30
0.03
0.33

143

Table B.2 – Knee Moment Data for Subject F2
Subject F2
Shoe
Trial #
Neutral
Neutral
Neutral
Lateral Stiff
Lateral Stiff
Lateral Stiff
T-Form 1
T-Form 1
T-Form 1
T-Form 2
T-Form 2
T-Form 2
T-Form 3
T-Form 3
T-Form 3
T-Form 4
T-Form 4
T-Form 4
T-Form 5
T-Form 5
T-Form 5

1
2
3
1
2
3
1
2
4
1
2
3
1
2
3
1
2
3
1
2
3

Knee Adduction/Abduction Moment
Normalized Joint Moment (Nm/kg) Average Standard Deviation
0.40
0.06
0.35
0.36
0.29
0.43
0.04
0.46
0.46
0.50
0.40
0.39
0.03
0.38
0.35
0.36
0.35
0.02
0.35
0.33
0.30
0.03
0.29
0.31
0.25
0.45
0.06
0.38
0.33
0.35
0.16
0.09
0.25
0.27
0.33

144

Table B.3 – Knee Moment Data for Subject F3
Subject F3
Shoe
Trial #
Neutral
Neutral
Neutral
Lateral Stiff
Lateral Stiff
Lateral Stiff
T-Form 1
T-Form 1
T-Form 1
T-Form 2
T-Form 2
T-Form 2
T-Form 3
T-Form 3
T-Form 3
T-Form 4
T-Form 4
T-Form 4
T-Form 5
T-Form 5
T-Form 5

1
2
3
1
2
3
1
2
3
1
2
3
1
2
3
1
2
3
1
2
3

Knee Adduction/Abduction Moment
Normalized Joint Moment (Nm/kg) Average Standard Deviation
0.42
0.45
0.03
0.47
0.46
0.45
0.47
0.06
0.54
0.43
0.51
0.48
0.49
0.04
0.43
0.49
0.45
0.04
0.45
0.42
0.40
0.44
0.04
0.48
0.45
0.44
0.44
0.03
0.46
0.41
0.48
0.46
0.03
0.43
0.46

145

Table B.4 – Knee Moment Data for Subject F4
Subject F4
Shoe
Trial #
Neutral
Neutral
Neutral
Lateral Stiff
Lateral Stiff
Lateral Stiff
T-Form 1
T-Form 1
T-Form 1
T-Form 2
T-Form 2
T-Form 2
T-Form 3
T-Form 3
T-Form 3
T-Form 4
T-Form 4
T-Form 4
T-Form 5
T-Form 5
T-Form 5

1
2
3
1
2
3
1
2
3
1
2
3
1
2
3
1
2
3
1
2
3

Knee Adduction/Abduction Moment
Normalized Joint Moment (Nm/k g) Average Standard Deviation
0.35
0.01
0.36
0.37
0.37
0.39
0.01
0.38
0.38
0.37
0.41
0.45
0.06
0.40
0.34
0.47
0.41
0.03
0.44
0.45
0.51
0.04
0.48
0.50
0.43
0.42
0.03
0.39
0.39
0.37
0.42
0.02
0.44
0.46
0.44

146

Table B.5 – Knee Moment Data for Subject M1
Subject M1
Shoe
Trial #
Neutral
Neutral
Neutral
Lateral Stiff
Lateral Stiff
Lateral Stiff
T-Form 1
T-Form 1
T-Form 1
T-Form 2
T-Form 2
T-Form 2
T-Form 3
T-Form 3
T-Form 3
T-Form 4
T-Form 4
T-Form 4
T-Form 5
T-Form 5
T-Form 5

1
2
3
2
3
4
2
4
5
1
2
3
1
2
3
1
2
3
2
3
4

Knee Adduction/Abduction Moment
Normalized Joint Moment (Nm/kg) Average Standard Deviation
0.72
0.75
0.04
0.79
0.73
0.50
0.56
0.57
0.08
0.66
0.71
0.73
0.73
0.03
0.76
0.76
0.81
0.76
0.05
0.71
0.86
0.81
0.06
0.74
0.83
0.74
0.75
0.07
0.83
0.69
0.78
0.82
0.80
0.02
0.81

147

Table B.6 – Knee Moment Data for Subject M2
Subject M2
Shoe
Trial #
Neutral
Neutral
Lateral Stiff
Lateral Stiff
T-Form 1
T-Form 1
T-Form 1
T-Form 2
T-Form 2
T-Form 3
T-Form 3
T-Form 3
T-Form 4
T-Form 4
T-Form 4
T-Form 5
T-Form 5
T-Form 5

1
2
1
3
1
2
3
2
3
1
2
3
1
2
3
1
2
3

Knee Adduction/Abduction Moment
Normalized Joint Moment (Nm/k g) Average Standard Deviation
0.42
0.37
0.08
0.31
0.28
0.28
0.00
0.28
0.28
0.27
0.08
0.19
0.34
0.16
0.21
0.07
0.26
0.18
0.22
0.06
0.28
0.19
0.28
0.28
0.01
0.27
0.29
0.20
0.25
0.11
0.18
0.38

148

Table B.7 – Knee Moment Data for Subject M3
Subject M3
Shoe
Trial #
Neutral
Neutral
Neutral
Lateral Stiff
Lateral Stiff
Lateral Stiff
T-Form 1
T-Form 1
T-Form 1
T-Form 2
T-Form 2
T-Form 2
T-Form 3
T-Form 3
T-Form 3
T-Form 4
T-Form 4
T-Form 4
T-Form 5
T-Form 5
T-Form 5

1
2
3
1
2
3
1
2
3
1
2
3
1
2
4
1
2
5
1
2
3

Knee Adduction/Abduction Moment
Normalized Joint Moment (Nm/kg) Average Standard Deviation
0.13
0.17
0.03
0.19
0.18
0.22
0.24
0.02
0.26
0.24
0.25
0.26
0.25
0.01
0.24
0.21
0.20
0.20
0.02
0.18
0.26
0.24
0.02
0.23
0.23
0.30
0.29
0.01
0.29
0.29
0.26
0.27
0.03
0.30
0.25

149

Table B.8 – Knee Moment Data for Subject M4
Subject M4
Shoe
Trial #
Neutral
Neutral
Neutral
Lateral Stiff
Lateral Stiff
Lateral Stiff
T-Form 1
T-Form 1
T-Form 1
T-Form 2
T-Form 2
T-Form 2
T-Form 3
T-Form 3
T-Form 3
T-Form 4
T-Form 4
T-Form 4
T-Form 5
T-Form 5
T-Form 5

1
2
3
1
2
3
1
2
3
1
2
3
1
2
3
1
2
3
1
2
3

Knee Adduction/Abduction Moment
Normalized Joint Moment (Nm/kg) Average Standard Deviation
0.30
0.38
0.07
0.43
0.42
0.28
0.31
0.03
0.32
0.34
0.32
0.38
0.37
0.05
0.42
0.35
0.41
0.39
0.03
0.41
0.31
0.35
0.03
0.36
0.37
0.37
0.35
0.03
0.32
0.37
0.34
0.35
0.02
0.37
0.35

150

APPENDIX C
BODYBUIDLER CODE FOR CHAPTER 3

151

Model Parameter File (.mp)
(Used for all models)
$HipOffset and $BodyMass were manually measured and input into the file
$KneeOffset and $Ankle Offset were measured and input by processing the static trial
{*T-Form.mp*}
{*Written by Jerrod Braman*}
{*For use with T-Form.mod and T-Form.mkr*}
{*All Distance measurements in millimeters*}
{*All angles in degrees*}
{*All mass in kilograms*}
{*General Parameters*}
{*==================*}
$HipOffset =
$MarkerDiameter = 14
$BodyMass =

{*Static Trial Parameters*}
$KneeOffset =
$AnkleOffset =
Marker Set File (.mkr)
(Used for all models)
!MKR#2
[Markers]
ASIS
RGTR
RPOT
RMKJ
RLKJ
RPOC
RMMA
RLMA
RLCA
R5MH
R2MH

Anterior Superior Iliac Spine
Right Greater Trochanter
Right Posterior Thigh
Right Medial Knee Joint Line
Right Lateral Knee Joint Line
Right Posterior Calf
Right Medial Malleolus
Right Lateral Malleolus
Right Lateral Calcaneous
Right 5th Metatarsal Head

152

Force1
Plate1
CenterPlate
RCKJ
RANK
RGTRHead

Force Value of Forceplate
COP Value
Center of the Force Plate
Right Center of the Knee Joint
Right Ankle Center
Right Greater Trochanter Head

R2MH,RANK
RANK,RCKJ
RCKJ,RGTRHead
Force1,Plate1
[Segments]

[Force Vectors]
P_ForcePlate1 Base of Plate 1 Vector
F_ForcePlate1 Tip of Plate 1 Vector
[Angles]
RKneeAngles Right Knee Rotation
RAnkleAngles Right Ankle Rotation
[Forces]
RKneeForce Right Knee Resultant Force
RAnkleForce Right Ankle Resultant Force
[Moments]
RKneeMoment Right Knee Resultant Moment
RAnkleMoment Right Ankle Resultant Moment

Massless Model
axislength = 100
{*Start of macro section*}

153

{*======================*}
Macro REPLACE4(p1,p2,p3,p4)
{*Replaces any point missing from set of four fixed in a
segment*}
s234 = [p3,p2-p3,p3-p4]

{*Defines a segment s234 using

all points except p1*}
p1V = Average(p1/s234)*s234

{*Finds the average position

of p1 in the s234 local Co-ord system and creates virtual
point p1V from this reference system*}
s341 = [p4,p3-p4,p4-p1]
{*Defines a segment s341 using
all points except p2*}
p2V = Average(p2/s341)*s341

{*Finds the average position

of p2 in the s341 local Co-ord system and creates virtual
point p2V from this reference system*}
s412 = [p1,p4-p1,p1-p2]
{*Defines a segment s412 using
all points except p3*}
p3V = Average(p3/s412)*s412

{*Finds the average position

of p3 in the s412 local Co-ord system and creates virtual
point p3V from this reference system*}
s123 = [p2,p1-p2,p2-p3]
{*Defines a segment s123 using
all points except p4*}
p4V = Average(p4/s123)*s123

{*Finds the average position

of p4 in the s123 local Co-ord system and creates virtual
point p4V from this reference system*}
{* Now only replaces if original is missing 11-99 *}
p1 = p1 ? p1V
p2 = p2 ? p2V
p3 = p3 ? p3V
p4 = p4 ? p4V
endmacro
Macro Axes(segment,axislength)
{* This macro creates segment axes for display purposes*}

154

segment#o={0,0,0}*segment
segment#x={axislength,0,0}*segment
segment#y={0,axislength,0}*segment
segment#z={0,0,axislength}*segment
output (segment#o,segment#x,segment#y,segment#z)
endmacro
{*Forceplate Data*}
{*===============*}
if EXIST( ForcePlate1 )
Force1 = ForcePlate1(1)
Moment1 = ForcePlate1(2)
Centre1 = ForcePlate1(3)
if ( ABS ( Force1 ) > 10 )
Point1 = Centre1 + {-Moment1(2)/Force1(3),
Moment1(1)/Force1(3), -Centre1(3) }
else
Point1 = Centre1
endif
Force1 = Force1 + Point1
OUTPUT ( Point1, Force1, Centre1 , Moment1 )
endif
{*Define Optional Points*}
{*======================*}
OptionalPoints(RMKJ,RMMA)
{*Define Offset Points*}
{*====================*}
RANK = RLMA+{0,$AnkleOffset,0}
RCKJ = RLKJ+{0,$KneeOffset,0}
RGTRHead = RGTR+{0,$HipOffset,0}
Output(RANK,RCKJ,RGTRHead)
If $Static == 1 Then
$KneeOffset = DIST(RLKJ,RMKJ)/2
$AnkleOffset = DIST(RLMA,RMMA)/2
Param($KneeOffset)
Param($AnkleOffset)
EndIf

155

{*Define Segments*}
{*===============*}
{*Foot Segment*}
{*============*}
RFoot = [R5MH,R5MH-RLCA,RCKJ-RANK,xyz]
{*Knee Segment*}
{*============*}
RShank = [RANK,RANK-RLMA,RCKJ-RANK,yxz]
{*Thigh Segment*}
{*============*}
RThigh = [RCKJ,RCKJ-RLKJ,RCKJ-RGTRHead,yxz]
Axes(RThigh,100)
Axes(RShank,100)
Axes(RFoot,100)
{*Anthropometric Data*}
AnthropometricData
RFootAnthro .0
.0
RShankAnthro
.0
RThighAnthro .0
.0
EndAnthropometricData

.0
.0
.0

0
.0
0

0

{*Kinetic Hierarchy and Associated Data*}
RThigh=[RThigh,RThighAnthro]
RShank=[RShank,RThigh,RCKJ,RShankAnthro]
RFoot=[RFoot,RShank,RANK,RFootAnthro]
AF=REACTION(RFoot)
KF=REACTION(RShank)
OUTPUT(AF,KF)
GAF=AF*RFoot
GKF=KF*RShank
AnkleForce=GAF(1)
AnkleMoment=GAF(2)
KneeForce=GKF(1)
KneeMoment=GKF(2)
OUTPUT(AnkleForce,AnkleMoment,KneeForce,KneeMoment)
Quasi-static Model

156

axislength = 100
{*Start of macro section*}
{*======================*}
Macro REPLACE4(p1,p2,p3,p4)
{*Replaces any point missing from set of four fixed in a segment*}
s234 = [p3,p2-p3,p3-p4]
{*Defines a segment s234 using all points except p1*}
p1V = Average(p1/s234)*s234
{*Finds the average position of p1 in the s234 local Co-ord
system and creates virtual point p1V from this reference system*}
s341 = [p4,p3-p4,p4-p1]
{*Defines a segment s341 using all points except p2*}
p2V = Average(p2/s341)*s341
{*Finds the average position of p2 in the s341 local Co-ord
system and creates virtual point p2V from this reference system*}
s412 = [p1,p4-p1,p1-p2]
{*Defines a segment s412 using all points except p3*}
p3V = Average(p3/s412)*s412
{*Finds the average position of p3 in the s412 local Co-ord
system and creates virtual point p3V from this reference system*}
s123 = [p2,p1-p2,p2-p3]
{*Defines a segment s123 using all points except p4*}
p4V = Average(p4/s123)*s123
{*Finds the average position of p4 in the s123 local Co-ord
system and creates virtual point p4V from this reference system*}
{* Now only replaces if original is missing 11-99 *}
p1 = p1 ? p1V
p2 = p2 ? p2V
p3 = p3 ? p3V
p4 = p4 ? p4V
endmacro
Macro Axes(segment,axislength)
{* This macro creates segment axes for display purposes*}
segment#o={0,0,0}*segment
segment#x={axislength,0,0}*segment
segment#y={0,axislength,0}*segment
segment#z={0,0,axislength}*segment
output (segment#o,segment#x,segment#y,segment#z)
endmacro
{*Forceplate Data*}
{*===============*}
if EXIST( ForcePlate1 )
Force1 = ForcePlate1(1)
Moment1 = ForcePlate1(2)
Centre1 = ForcePlate1(3)
if ( ABS ( Force1 ) > 10 )
Point1 = Centre1 + {-Moment1(2)/Force1(3),

157

Moment1(1)/Force1(3), -Centre1(3) }
else
Point1 = Centre1
endif
Force1 = Force1 + Point1
OUTPUT ( Point1, Force1, Centre1 , Moment1 )
endif
{*Define Optional Points*}
{*======================*}
OptionalPoints(RMKJ,RMMA)
{*Define Offset Points*}
{*====================*}
RANK = RLMA+{0,$AnkleOffset,0}
RCKJ = RLKJ+{0,$KneeOffset,0}
R2MH ={R5MH(1),RANK(2),R5MH(3)}
RGTRHead = RGTR+{0,$HipOffset,0}
{*COM for segments*}
comf=(RANK+R2MH)/2
coms=RANK+(RCKJ-RANK)*.567
Output(RANK,RCKJ,RGTRHead,comf,coms)
If $Static == 1 Then
$KneeOffset = DIST(RLKJ,RMKJ)/2
$AnkleOffset = DIST(RLMA,RMMA)/2
Param($KneeOffset)
Param($AnkleOffset)
EndIf
{*Define Segments*}
{*===============*}
{*Foot Segment*}
{*============*}
RFoot = [R5MH,R5MH-RLCA,RCKJ-RANK,xyz]
{*Knee Segment*}
{*============*}
RShank = [RANK,RANK-RLMA,RCKJ-RANK,yxz]
{*Thigh Segment*}

158

{*============*}
RThigh = [RCKJ,RCKJ-RLKJ,RCKJ-RGTRHead,yxz]
Axes(RThigh,100)
Axes(RShank,100)
Axes(RFoot,100)
{*Anthropometric Data*}
AnthropometricData
RFootAnthro .0
.0
RShankAnthro
.0
RThighAnthro .0
.0
EndAnthropometricData

.0
.0
.0

0
.0
0

0

{*Kinetic Hierarchy and Associated Data*}
footforce=$BodyMass*.0146*9.8
dummyforce={0,0,-footforce}
dummymoment={0,0,0}
dummyapplication={comf(1),comf(2),comf(3)}
reactionforce=|dummyforce,dummymoment,dummyapplication|
LRF=reactionforce/RFOOT
CONNECT(RFOOT,reactionforce,1)
shankforce=$BodyMass*.0465*9.8
dummyforce2={0,0,-shankforce}
dummymoment2={0,0,0}
dummyapplication2={coms(1),coms(2),coms(3)}
reactionforce2=|dummyforce2,dummymoment2,dummyapplication2|
CONNECT(RSHANK,reactionforce2,1)
RThigh=[RThigh,RThighAnthro]
RShank=[RShank,RThigh,RCKJ,RShankAnthro]
RFoot=[RFoot,RShank,RANK,RFootAnthro]
AF=REACTION(RFoot)
KF=REACTION(RShank)
OUTPUT(AF,KF)
GAF=AF*RFoot
GKF=KF*RShank
AnkleForce=GAF(1)
AnkleMoment=GAF(2)

159

KneeForce=GKF(1)
KneeMoment=GKF(2)
OUTPUT(AnkleForce,AnkleMoment,KneeForce,KneeMoment)

Full Dynamic Model
axislength = 100
{*Start of macro section*}
{*======================*}
Macro REPLACE4(p1,p2,p3,p4)
{*Replaces any point missing from set of four fixed in a
segment*}
s234 = [p3,p2-p3,p3-p4]

{*Defines a segment s234 using

all points except p1*}
p1V = Average(p1/s234)*s234

{*Finds the average position

of p1 in the s234 local Co-ord system and creates virtual
point p1V from this reference system*}
s341 = [p4,p3-p4,p4-p1]
{*Defines a segment s341 using
all points except p2*}
p2V = Average(p2/s341)*s341

{*Finds the average position

of p2 in the s341 local Co-ord system and creates virtual
point p2V from this reference system*}
s412 = [p1,p4-p1,p1-p2]
{*Defines a segment s412 using
all points except p3*}
p3V = Average(p3/s412)*s412

{*Finds the average position

of p3 in the s412 local Co-ord system and creates virtual
point p3V from this reference system*}
s123 = [p2,p1-p2,p2-p3]
{*Defines a segment s123 using
all points except p4*}
p4V = Average(p4/s123)*s123

{*Finds the average position

of p4 in the s123 local Co-ord system and creates virtual

160

point p4V from this reference system*}
{* Now only replaces if original is missing 11-99 *}
p1 = p1 ? p1V
p2 = p2 ? p2V
p3 = p3 ? p3V
p4 = p4 ? p4V
endmacro
Macro Axes(segment,axislength)
{* This macro creates segment axes for display purposes*}
segment#o={0,0,0}*segment
segment#x={axislength,0,0}*segment
segment#y={0,axislength,0}*segment
segment#z={0,0,axislength}*segment
output (segment#o,segment#x,segment#y,segment#z)
endmacro
{*Forceplate Data*}
{*===============*}
if EXIST( ForcePlate1 )
Force1 = ForcePlate1(1)
Moment1 = ForcePlate1(2)
Centre1 = ForcePlate1(3)
if ( ABS ( Force1 ) > 10 )
Point1 = Centre1 + {-Moment1(2)/Force1(3),
Moment1(1)/Force1(3), -Centre1(3) }
else
Point1 = Centre1
endif
Force1 = Force1 + Point1
OUTPUT ( Point1, Force1, Centre1 , Moment1 )
endif
{*Define Optional Points*}
{*======================*}
OptionalPoints(RMKJ,RMMA)
{*Define Offset Points*}
{*====================*}
RANK = RLMA+{0,$AnkleOffset,0}
RCKJ = RLKJ+{0,$KneeOffset,0}

161

R2MH ={R5MH(1),RANK(2),R5MH(3)}
RGTRHead = RGTR+{0,$HipOffset,0}
{*COM for segments*}
comf=(RANK+R2MH)/2
coms=RANK+(RCKJ-RANK)*.567
Output(RANK,RCKJ,RGTRHead,R2MH,comf,coms)
If $Static == 1 Then
$KneeOffset = DIST(RLKJ,RMKJ)/2
$AnkleOffset = DIST(RLMA,RMMA)/2
Param($KneeOffset)
Param($AnkleOffset)
EndIf
{*Define Segments*}
{*===============*}
{*Foot Segment*}
{*============*}
RFoot = [R5MH,R5MH-RLCA,RCKJ-RANK,xyz]
{*Knee Segment*}
{*============*}
RShank = [RANK,RANK-RLMA,RCKJ-RANK,yxz]
{*Thigh Segment*}
{*============*}
RThigh = [RCKJ,RCKJ-RLKJ,RCKJ-RGTRHead,yxz]
Axes(RThigh,100)
Axes(RShank,100)
Axes(RFoot,100)
{*Anthropometric Data*}
AnthropometricData
RFootAnthro .0145 .50
.475
RShankAnthro
.0465 .567
RThighAnthro .100 .567 .323
EndAnthropometricData

0
.302
0

0

{*Kinetic Hierarchy and Associated Data*}

162

RThigh=[RThigh,RThighAnthro]
RShank=[RShank,RThigh,RCKJ,RShankAnthro]
RFoot=[RFoot,RShank,RANK,RFootAnthro]
AF=REACTION(RFoot)
KF=REACTION(RShank)
OUTPUT(AF,KF)
GAF=AF*RFoot
GKF=KF*RShank
AnkleMoment=GAF(2)
KneeMoment=GKF(2)
OUTPUT(AnkleMoment,KneeMoment)

163

APPENDIX D
BODYBUILDER CODE FOR CHAPTER 4

164

Model Parameter File (.mp)
{*WaveDisk.mp*}
{*Written by Jerrod Braman*}
{*For use with WaveDisk.mod and WaveDisk.mkr*}
{*All Distance measurements in millimeters*}
{*All angles in degrees*}
{*All mass in kilograms*}
{*General Parameters*}
{*==================*}
Marker Set File (.mkr)
!MKR#2
[Markers]
Heel1 Top Heel
Heel2 Center Heel
Heel3 Right Heel
Toe1 Top Toe
Toe2 Center Toe
Toe3 Right Toe
Heel1,Heel2
Heel2,Heel3
Toe1,Toe2
Toe2,Toe3
Biomechanical Model File (.mod)
{*WaveDisk Analysis Model*}
{*Orthopaedic Biomechanics Laboratories -- Gait Analysis Laboratory*}
{*Written by Jerrod Braman*}
{*Use with BodyBuilder*}
{*Use with WaveDisk.mkr*}
{*Use with WaveDisk.mp*}
axislength = 100
Macro Axes(segment,axislength)

165

{* This macro creates segment axes for display purposes*}
segment#o={0,0,0}*segment
segment#x={axislength,0,0}*segment
segment#y={0,axislength,0}*segment
segment#z={0,0,axislength}*segment
output (segment#o,segment#x,segment#y,segment#z)
endmacro
{*Define Segments*}
{*===============*}
{*Heel Segment*}
{*============*}
Heel = [Heel2,Heel3-Heel2,Heel1-Heel2,yxz]
{*Toe Segment*}
{*============*}
Toe = [Toe2,Toe3-Toe2,Toe1-Toe2,yxz]
{*Global Segment*}
GlobalO={0,0,0}
Globalx={100,0,0}
Globaly={0,100,0}
Global=[GlobalO,Globalx-GlobalO,GlobalO-Globaly,xzy]
OUTPUT(GlobalO,Globalx,Globaly)
Axes(Heel,100)
Axes(Toe,100)
Axes(Global,100)
{*Angle Calculations*}
ToeAngle=<Toe,Global,xyz>
HeelAngle=<Heel,Global,xyz>
OUTPUT(ToeAngle,HeelAngle)

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APPENDIX F
OTHER BIOMECHANICAL PROJECTS

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Ph.D. Project
Biomechanical comparison of three methods of back squatting
Adam Bruenger, Ph.D. – 2009

M.S. Projects
An analysis of pressure distribution with a prefabricated foot orthotic on a symptomatic
population
B.J. Vascik, M.S. – 2006
The difference in immediate changes in dorsiflexion range of motion using an ultrasound heat
treatment, followed by two different stretching techniques
Gregory Hawthorne, Jr., M.S., ATC – 2008
A comparison of the effects of joint mobilizations versus muscle energy on increasing shoulder
range of motion in healthy individuals
Anna Leyland, M.S. – 2009
Assisted in designing and manufacturing data collection device and data collection

Wolverine World Wide, Incorporated Projects
Parotec In-shoe Pressure Measurement System
Thermal Footwear Characteristics (Subject and Mechanical Testing)
Compression Testing (Instron)
Impact Testing (ASTM Drop Test)
Flexion Testing
Force Plate Testing
Vicon Motion Analysis

Wound Healing Shoe Project
Evaluated a would healing shoe using the Parotec In-shoe Pressure Measurement System

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