A study of plantar stresses underneath metatarsal heads in the human foot 4

92








CHAPTER 4
EVALUATION OF MODEL
PERFORMANCE



~ Success is not measured by what you accomplish, but by the opposition
you have encountered, and the courage with which you have maintained
the struggle against overwhelming odds~
Orison Swett Marden


93
4.1 Comparison of FEM results with experimental data
Finite element analyses and experimental studies (in vivo study and in vitro
cadaveric testing) are complimentary techniques to characterize the complex
biomechanical behavior of the foot and its components, involving muscular
control, joint movements, bone stresses, as well as the foot-ground interactions.
In general, experimental investigations have provided clinically relevant data (e.g.
plantar pressure distribution, foot deformities etc.) with regards to normal and
pathological foot conditions. However, in terms of plantar shear stresses, internal

tissue stresses, and roles of muscles, ligaments, plantar fascia, and other soft
tissue structures in the weight-bearing function of the foot, these experimental
studies are often limited. The FE models of the foot thus supplement those
experimental studies. Once such models are verified by experimental data, they
offer additional extrapolated information.
In this study, a comprehensive finite element musculoskeletal model of the
foot was constructed. Compared to the majority of existing foot models which
focus on standing (Gefen, 2002, Cheung et al., 2005, Chen et al., 2010a), the
current model accommodates realistic musculoskeletal loads and simulates a
muscle-demanding posture corresponding to heel-rise, which facilitates a
detailed investigation into the foot mechanism that involves complex interplay of
muscular control, articulating joint movements, and forefoot plantar loading
distributions. Meanwhile, great efforts were made experimentally in order to 1)
provide a model with accurate material property inputs for the sub-MTH soft
tissue; 2) validate the model’s predictions for foot-ground interactions at local

94
MTH sites. In this section, an experimental validation will be conducted by
comparing of FEM results with experimental data obtained from the developed
gait platform and those reported in the literature. It will be demonstrated that with
accurately quantified tissue property, muscular loading characteristics, and foot’s
geometric positioning, a realistic stress response of the model during foot-ground
interactions can be reproduced.

4.1.1 Simulated muscular loads
Since the number of flexor muscles present in the model is larger than
necessary to achieve mechanical equilibrium for all foot geometrical positioning,
the system is indeterminate. In other words, infinitely many different muscle force
combinations can balance the same foot geometrical positioning during stance.
Thus, for simplification, one of the most important assumptions made in the

modeling study is for the calculation of muscle forces, i.e. the extrinsic flexor
tendons were loaded to their relative strengths according to the physiological
cross-sectional areas (PCSA) of their respective muscle bundle. However, this
assumption may be justified by the fact that the force-generating capacity of a
muscle is known to be directly proportional to its PCSA (Wickiewicz et al., 1983,
Fukunaga et al., 1996). With this assumption, it possible to inversely deduce the
combinations of extrinsic flexor forces required to generate the targeted GRF that
match the given boundary conditions.
When comparing the FE model calculated muscular loads with literature
data, the calculated G-S muscle force of 1620N is very close to the in vivo peak

95
Achilles tendon loading (1430 ±500 N) measured by Finni et al. (1998) for
walking. Forces in the other extrinsic muscles were also within their physiological
ranges and generally comparable to those utilized in actuating a cadaveric foot
(Sharkey and Hamel, 1998) to emulate walking. A detailed quantitative
comparison of the calculated six flexors muscle forces used in the current finite
element model with those estimated by Salathe and Arangio (2002) in an
analytical foot structure model is presented in Fig. 4.1.1.1.


Fig. 4.1.1.1. Comparison of the muscular loads obtained from the finite element model
with those calculated by Salathe and Arangio (2002) and Gefen (2000)

In the analytical model proposed by Salathe and Arangio (2002), the
anatomical structures of the foot bones, ligaments, and muscles were considered
with foot bones being articulated through various joints, such as hinge and
universal joints. Due to the indeterminacy encountered for calculating the muscle
forces, the authors made a similar assumption that each muscle of a given group

96
contributes to the total support provided by that group an amount proportional to
its PCSA. Good agreement between the numerical and analytical values of flexor
muscle forces was found. The muscular loads applied in the current model were
also quantitatively similar to those calculated by Gefen et al. (2000) in a 3-D FE
foot model simulating stance-phase of walking. Thus, the muscles modeled and
magnitudes of forces simulated are considered representative of a specific
instant in the gait cycle, corresponding approximately to the occurrence of the
second GRF peak during walking.

4.1.2 Plantar stress distribution
With all muscular loads applied, the foot model was successfully solved in a
typical geometry following heel rise, with the ankle and MTP joints maintained at
plantar-flexed and extended configurations, respectively. Corresponding to this
foot posture, the predicted contours of plantar pressure and shear stress
distributions were plotted in Fig. 4.1.2.1. The plantar pressure was mainly
concentrated at areas under the 2nd and 3rd MTHs, with peak plantar pressure
of 570.6 KPa found under the 2nd MTH. This is generally in agreement with
literature data on foot roll-over characteristics, i.e. after heel off, the forefoot has
a more central push-off over the second metatarsal at the terminal stance of
walking (De Cock et al., 2005). While for the predicted plantar shear, it shows a
different stress patterns from the plantar pressure. The peak anterior-posterior
(AP) shear stress of 123.6 kPa was found under the 1st MTH and the peak
medial-lateral (ML) shear stress of 55.7 kPa was found under the 2nd MTH. The

97
peak shear stresses found were generally much lower than the peak plantar
pressure values.

Fig. 4.1.2.1. Model predicted (A) plantar pressure, (B) plantar shear (vector summation
of AP and ML components), and (C) tissue VMS stress distributions

To facilitate direct comparison of the plantar stresses calculated by the
model and measured by the sensor array, the continuous predicted plantar
pressure data were averaged over the sensing area of an individual sensor (9 ｘ
9 mm
2
). When compared to the averaged plantar stress data, the model agreed
extremely well in peak plantar pressure values (570.6 kPa Versus 568.2 kPa).
For peak AP and ML shear stresses, the peak shear location compared relative
well with experimental data. The peak shear predicted by the model was higher
than what was observed experimentally (123.6 kPa Versus 74.2 kPa) (Table
4.1.2.1.).


98
Table 4.1.2.1. Summary of peak plantar stresses measured by sensor array and
predicted by the finite element model under individual metatarsal heads (MTHs) (PPP:
peak plantar pressure; PSS-AP: peak anterior-posterior shear stress; PSS-ML: peak
medial-lateral shear stress. Unit: kPa)
Sensor measured plantar
stresses
Model predicted contact
stresses
MTH
location
PPP PSS-AP PSS-ML PPP PSS-AP PSS-ML
1st 290.1 74.2 21.9 143.8 123.6 20.0
2nd 568.2 41.1 32.8 570.6 17.9 55.7

3rd 412.2 30.3 24.1 513.7 15.1 45.1
4th 313.9 50.1 8.1 267.2 36.1 9.9
5th 129.1 10.4 7.5 17.1 (<0.1) (<0.1)
Mean
(±Std)
342.7±130 41.2±23 18.9±10 302.5±236 48.2±51 32.7±21

Since the interfacial property (frictional contact at foot-ground interface)
used in this model was experimentally determined, the authors believe the
reason for inadequate prediction of shear is the fact that the vertical and the
shear force components may not have a simple linear relationship that follows
classic Coulomb's law of friction. Plantar shear magnitudes probably are
influenced by local differences in frictional properties and/ or intrinsic muscle
activities which were not accounted for the FE model. Nevertheless, the pattern
of both AP and ML shear stress distribution, in terms of peak shear found under
individual MTHs compare well with experimental data.




99

4.1.3 Metatarsal bone strain
The model was further substantiated by comparing predicted metatarsal
strains to results obtained from the cadaveric work of Sharkey et al., (1995). This
cadaver study had similar loading conditions with the current FE model. In their
experiments, each foot was loaded to a maximum of 750N of ground-reaction
force by simulated contraction of the triceps surae, and strains were recorded in
the mid-part of the shaft of the second metatarsal. For the current FE model,
local coordinate systems were defined for the metatarsal elements to obtain axial

surface strain components directly from FE analyses. Simulation results were
extracted from the mid-spans of the 2nd metatarsal shafts; nodal values with an
area of 5 ｘ 10 mm were averaged for comparison with strain gauge
measurements by Sharkey et al. (1995). The predicted strains were highest at
the dorsal mid-shaft region (2395.1με), followed by those at the medial area
(946.2με), and they were lowest at the lateral mid-shaft vicinity (376.7με); these
were comparable to experimental data (Table 4.1.3.1.), suggesting that the finite
element model provides a reasonable reflection of the internal foot mechanics,
and is thus a suitable tool for further investigation.
Table 4.1.3.1. Model predicted average strain in the 2nd metatarsal mid-shaft in
comparison with the cadaver study results by Sharkey et al., (1995).

Dorsal Strain (με) Medial Strain (με) Lateral Strain (με)
Sharkey et al.,
(1995)
-1700±884 733±1651 422±819
FE Model 2395.1 946.2 263.3


100

4.2 Sensitivity of model to flexor muscle force variations
The present model is a truly musculoskeletal FE model which allows quantitative
analysis of altered muscular control on biomechanical behavior of the foot. The
aim of this sensitivity analysis was to address a clinical-relevant topic: what role
does the gastrocnemius-soleus (G-S) complex plays in normal weight-bearing
function by the foot. The results obtained from this model’s sensitivity analysis
may have important clinical implications on tendo-Achilles lengthening
procedures, and to provide surgeons with an understanding of the underlying
mechanism for relieving forefoot pressure in diabetic patients suffering from

ankle equinus contracture.

4.2.1 Role of the gastrocnemius-soleus (G-S) muscle complex
The functional roles of the extrinsic plantar flexor muscles are to provide stability
to the foot during the stance phase of gait (Sutherland et al., 1980). Anatomically,
these musculotendinous units usually possess unique long-tendon structures,
which allow them to influence multiple joints inside the foot. This multi-articular
characteristic in extrinsic flexors facilitate control and stabilization of major bony
joints, including the talo-crural, the subtalar, and indirectly the
metatarsophalangeal (MTP) joints, thus providing primary coordination of the
stance-phase placement of the foot, which is essential for normal weight-bearing
during gait (Perry, 1992).

101
The gastrocnemius-soleus (G-S) complex associated with the Achilles
tendon is the most dominant extrinsic plantar flexor. Various studies using
cadaveric foot models have established the important biomechanical linkage
among the Achilles tendon, plantar fascia and MTP joints, both statically (Carlson
et al., 2000) and dynamically (Erdemir et al., 2004), earlier described by Hicks
(1955), as the Windlass mechanism. With the foot stabilized by flexor muscles
upon heel-rise, substantial dynamic ground reaction forces (GRF) are imposed
only on the forefoot, thereby generating highly localized stresses, particularly
underneath the metatarsal heads (MTHs). Such potentially ‘detrimental’ stresses
may only last for a short-dwelling time in a normal foot, provided that the
coordinative muscles involved activate the right amount of force at the exact
appropriate time during gait (Perry, 1992, Hayafune et al., 1999).
When equinus contracture is present, normal muscle activation in the G-S
complex may be compromised (Damron et al., 1994, Lin et al., 1996). The
contractured G-S muscle consistently exerts tension through the Achilles tendon,
resulting in an earlier Windlass effect accompanied by premature heel-rise, and

the maintenance of this abnormal posture throughout the whole stance phase of
walking (D'Ambrogi et al., 2005, Hastings et al., 2000). Many believe that this
leads to prolonged and high-magnitude forefoot loading exposure, which may
induce pain, injury or other structural deformities in these areas over time, such
as metatarsalgia and plantar forefoot ulceration (Bowers and Castro, 2007).
Manipulation of the muscle force in the G-S complex via surgical
intervention, such as tendo-Achilles lengthening, can negate the ‘abnormal’

102
Windlass mechanism associated with G-S muscle contracture (Salsich et al.,
2005), and appears effective in restoring normal forefoot load transfer by
relieving the abnormally high stresses under the MTHs (Armstrong et al., 1999,
Maluf et al., 2004). This facilitates the healing of ulcers within that vicinity in a
diabetic foot (Lin et al., 1996, Mueller et al., 2003). In contrast, Orendurff et al.
(2006) showed that equinus contracture accounts for only a small amount of the
increased forefoot plantar pressure. A major point of debate has been the precise
role the G-S complex plays in normal weight-bearing by the foot; there is no
comprehensive documentation of this topic (McGuire, 2010).
Only few investigators have attempted experimental studies on the role of
extrinsic flexor muscles on forefoot skeletal loads and plantar stress distribution.
Most of these studies focus on the effects of one over another. The
interrelationship among muscular control, internal joint movements, and plantar
loading distributions have not been fully explored due to experimental difficulties.
Sharkey et al. (1995) simulated flexor muscle action in a cadaveric foot model
and successfully achieved a maximum ground reaction force of 750 N during
heel-lift; strains in the second metatarsal bone were also measured. Similar
loading conditions were applied by Ferris et al. (1995), in their study on the
influence of extrinsic plantar flexors on loads borne by the toes; no variation of
Achilles tendon forces was considered. In another in vitro investigation, Aronow
et al., (2006) found that upon heel-rise, the plantar pressure shifts from the

hindfoot to the forefoot when Achilles tendon forces are increased, but the
applied muscle loading was only 93 N which is much lower than physiological

103
values (e.g. according to Finni et al. (1998), the in-vivo measured peak Achilles
tendon force can range from 750 to 2360 N ).
Previous finite element (FE) modeling of the foot has demonstrated the
potential of such numerical modeling to investigate the effects of specific
changes to the foot structure on joint movements and stress responses, via
parametric analyses. The aims of the present model sensitivity study are to
quantitatively analyze adaptive changes of the foot mechanism, i.e. movements
at the ankle and metatarsophalageal joints and forefoot load transfer in response
to force variations generated by the G-S muscle complex. Precise determination
of such relationships may help explain the etiology of many foot and ankle
disorders associated with G-S contracture, and an understanding of the possible
benefits, if any, related to these biomechanical quantities following surgical
tendo-Achilles lengthening. This study also directly addresses, some significant
limitations of previous FE models related to foot biomechanics, namely the
absence of muscle actuation and stabilization.









104
4.2.2 Parametric study on F

AT
variations
To study the effects of varying the G-S complex muscle force (denoted as F
AT
)
on foot mechanism, a multi-step analysis procedure was adopted. The initially
loaded baseline model served as the reference for the subsequent analysis, in
which the maximum GRF was maintained whereby the maximum F
AT
was
reduced in steps of 10% down to 60% of the original, to simulate reduced muscle
effectiveness of the G-S complex (Table 4.2.2.1). The stresses and strains of the
baseline model were updated in sequential steps to elicit the results of
decreasing the F
AT
.

Table 4.2.2.1. Variations of G-S complex muscle force (F
AT
) studied in the present model
sensitivity analysis
Variations 100% 90% 80% 70% 60%
F
AT

1620N 1458N 1296N 1134N 972N


The movements at the ankle and metatarsophalangeal joints in the main
loading plane arising from adjustments in the muscle forces were analyzed.

Changes in the foot-ground interactions were characterized by the pressure
patterns displayed by contacting elements at the plantar surface. The data
obtained was compared with those of the baseline model.





105
4.2.2.1 Effects of Achilles tendon forces variations on movements at the ankle
and metatarsophalangeal (MTP) joints

Fig. 4.2.2.1. Correspondence of gastrocnemius-soleus muscle forces (indicated as F
AT
)
and the ankle () and metatarsophalangeal () joint angles during heel rise. The
amount of the muscle forces in F
AT
varied between 1620N (100% effectiveness) and
972N (60% effectiveness). The angles were calculated relative to the joint
configurations at foot’s neutral position.

With the full muscular load applied, the foot model corresponding to a
typical geometry at push-off, with the ankle and MTP joints maintained at
plantar-flexed and extended configurations, respectively, was examined. The
predicted joint angles were significantly affected by variations (i.e. reduction) in
F
AT
(Fig. 4.2.2.1.). With decreased muscle stabilization, the ankle joint rotates
by 8.81° from plantarflexion (6.05°) to dorsiflexion (-2.76°). This is coupled to

MTP joint movements with an average decrease from -9.04° to -4.39° in joint

106
extension.
The bony joint movements estimated in the parametric analysis
employing the current model are comparable to those observed in vivo clinically,
following surgical tendon-Achilles lengthening procedures. Previous studies
have documented a 9 ~19° increase in the passive ankle dorsiflexion range of
motion measured during non-weight-bearing conditions after such treatment
(Lin et al., 1996, Armstrong et al., 1999, Mueller et al., 2003). For loaded states,
the current model shows that the reduction in muscular tension via the Achilles
tendon can cause persistent increased dorsal-directed rotations of the ankle by
8.81°, and a decrease in MTP joint extension by 4.65°. This indicates the
central role of the gastrocnemius-soleus complex in modulating ankle mobility,
and its indirect effects on metatarsophalangeal joint configurations due to the
windlass mechanism, during heel rise (Hicks, 1955, Erdemir et al., 2004).
Interestingly, Maluf et al., (2004) recently reported that diabetic patients who
underwent tendon-Achilles lengthening experienced an average increase in
ankle dorsiflexion motion, ranging from 6.1 to 8.3°, measured during gait, which
appears more consistent with the present findings of 8.81°. In addition, the
plantar flexor moment reduction reported, ranging from 23% to 53% (Maluf et
al., 2004), encompasses the simulated F
AT
reduction in the current study. This
suggests a relationship between the muscle functional potential (i.e. torque) of
the G-S and the range of motion of the ankle joint in loaded states, as
demonstrated by the proposed model. The FEM results obtained may be
extrapolated to those in vivo conditions.

107

4.2.2.2 How Achilles tendon forces variations affects the forefoot plantar
pressure and shear stress distribution


Fig. 4.2.2.2. Changes in plantar pressure distributions for different angles at ankle
joints in response to incrementally reduced gastrocnemius-soleus (G-S) muscle forces.
Note that, as a result of altered foot geometry due to internal joint re-arrangements, the
initially concentrated pressure (under 2nd MTH) tends to redistribute the plantar loads
more equally among the five foot rays.

The effect of a decreased F
AT
on the forefoot stress distribution is
marginal (Fig. 4.2.2.2.). In terms of plantar pressure, the location of the highest
pressure predicted shifted in the distal-proximal direction, from the 2nd MTH to
the 1st, 3rd, and 4th MTHs. This was accompanied by an expansion of the
contact area from the metatarsal region to the distal side (i.e. close to mid-foot

108
region), resulting in a decrease of the overall peak plantar pressure from 321.1
to 261.7 kPa. In terms of tissue shear stress distributions, the peak plantar
shear stress was shirted from the 1
st
MTH site to the lateral MTH sites (Fig.
4.2.2.3.).


Fig. 4.2.2.3. Resultant plantar tissue shear stress redistribution in response to
incrementally reduced gastrocnemius-soleus (G-S) muscle forces.

By using pressure platform systems, Armstrong et al. (1999) & Maluf et
al. (2004), found that the peak pressure on the plantar aspect of the forefoot
was reduced significantly by 27% after tendon-Achilles lengthening in diabetic
patients who are at high risk of plantar ulceration. The present study shows that

109
the overall peak plantar pressure decreases by 22.7% in response to
prescribed reductions in the F
AT
. Although these changes appear associated
with increased ankle dorsiflexion, the actual cause could be the smaller
inclination angle of the metatarsals to the horizontal, leading to more even load
distribution among the five foot rays. Such plantar load redistribution could be
extremely sensitive, since the average angular changes at MTP joints were
only less than 5°. This is also evident in the parametric model, which showed
that plantar-directed displacements of the MTHs generally decreased by about
42% (Fig. 4.2.2.1.). A similar reason has also been suggested by D'Ambrogi et
al. (2005), that releasing the load on the Achilles tendon by plantar flexion of
the metatarsals, may alleviate the risk of cavoid foot formation, which frequently
occurs in diabetic patients. This again, highlights the importance of Achilles
tendon loading in forefoot weight-bearing, which is closely related to stance-
phase placement of the foot.

110
4.2.2.3 Effects of F
AT
variations on sub-MTH pressure peaks


Fig. 4.2.2.4. Percentage changes of local pressures under individual MTHs for different
ankle joint angles (). Nodal points just underneath five MTHs were selected for analysis.
Note the most significant pressure reduction site was under the 2nd MTH for the
particular model analyzed in this study.

Variations of local pressure peak at individual MTHs are plotted in Fig.
4.2.2.4. Note that there are site-specific pressure reductions and increases.
The most prominent relief of focal stresses was found underneath the 2nd and
3rd MTHs. Peak pressures there were reduced by 20.1% and 14.6%,
respectively. Peak pressure changes under the 4th MTH were negligible. There
was 6.9% increase in the 1st MTH peak pressure for the initial 20% reduction in
the F
AT
, but no further changes thereafter. The peak pressure at the 5th MTH
was almost doubled when F
AT
was reduced by 40%, although the absolute
stress values remain low. Since the current model in parametrical analysis was
implemented with a loading control regime. i.e. the net GRF remains

111
unchanged in all simulations, these would mean a more equal load distribution,

with the loads transferred from the central MTH region to the lateral MTHs. The
trends noted may have important clinical implications, in that areas under MTHs
other than the central ones (2nd and 3rd MTHs), are less likely to benefit from
tendon-Achilles lengthening procedures (i.e. relief of high pressures). However,
it should be noted that the pressure redistribution patterns obtained could
probably have arisen from structural alignment of the five MTHs in the
transverse plane, which is known to exhibit a unique hyperbola pattern in a
normal foot (Demp, 1998), as is the case for the particular model in this study.
The current model did not account for any deformed metatarsal structures
which may be present in the feet of patients who suffer from contracture of the
G-S complex.
In summary, the current model’s sensitivity analysis has predicted
alterations in the biomechanical behavior of the foot mechanism resulting from
variations of muscle strength in the G-S complex. The stance-phase placement
of the foot was found to be extremely sensitive to changes in such muscle loads,
which leads to forefoot plantar pressure and shear stress redistribution. The
findings of interrelationships among muscular control, internal joint movements,
and plantar loading distribution are envisaged to have important clinical
implications on tendo-Achilles lengthening procedures, and also provide
surgeons with knowledge to understand the underlying mechanism for relieving
forefoot pressure in diabetic patients who suffer from ankle equinus contracture.


112
4.3 Possible application of model in design of therapeutic footwear
Therapeutic footwear plays an important role in the prevention of neuropathic
plantar ulceration in DM patients, as it is a first line of defense against abnormal
plantar loading transfer (Caputo et al., 1994, Cavanagh et al., 1993b). When
plantar stress relief is the goal, optimal design of therapeutic footwear is often
required. Computational modeling of foot and footwear offers a unique

opportunity to conduct such design optimization, as it allows multiple design
variables of footwear to be ‘virtually’ tested in a more controlled and efficient
manner than traditional experiments (Lemmon et al., 1997, Chen et al., 2003,
Cheung et al., 2005, Erdemir et al., 2005, Actis et al., 2008). In this section,
possible application of the developed foot FE model in therapeutic footwear
design was demonstrated. More specifically, effects of modified foot-supporting
interface by using metatarsal support on plantar stress redistribution underneath
the MTHs were studied.

4.3.1 Insoles for tissue stress relief under MTHs
Within the plantar metatarsal region lie metatarsal heads (MTHs),
metatarsophlageal (MTP) joints, and the underlying protective fat pads, which are
subjected to complex tissue loading of pressure (i.e., arising from vertical force)
and shear stresses (i.e., arising from horizontal forces), when attenuating the
highest ground reaction forces (GRFs). Such focal stresses, that the forefoot
encountered during the terminal stance-phase of human gait, may be risk factors
for developing a number of musculoskeletal disorders in foot, such as pain in the

113
MTHs (e.g. metatarsalgia in rheumatoid arthritis) (Roy, 1988), joint structural
deformities (e.g. hallux valgus) (Sanders et al., 1992), and skin lesions in the
plantar soft tissue (e.g. diabetic foot ulcers) (Cavanagh et al., 1993). Many
investigators believed that shear stress can be more detrimental than pressure in
causing tissue injuries in the insensate diabetic foot.
The use of appropriate orthotic devices to relieve focal stresses in the
forefoot may aid in the treatment and prevention of these conditions, although
most are primarily focused on peak plantar pressure relief. For overall pressure
reduction, conservative strategies commonly involve prescription of extra-depth
orthopedic shoes with soft insoles (Ashry et al., 1997, Tsung et al., 2004), and
rocker bottom shoes with custom molded insert based on foot shape (Owings et

al., 2008), to facilitate the distribution of loads to the entire plantar surface. While
for extra pressure-relief at specific local MTHs, other shoe modifications, such as
inclusion of more ‘compliant’ midsole plugs into the footwear design, although
this may cause unwanted stress concentrations at plug edges, due to boundary
material discontinuity (Erdemir et al., 2005). Alternatively, metatarsal pads may
be used without these side-effects, and have potential to alter plantar loads (both
vertical and shear forces) transmission through the MTHs. These components
are in teardrop shape with a convex dome, and have traditionally been positioned
proximal to the target MTH (e.g. 5 mm) (Hastings et al., 2007). In this manner, a
fulcrum effect may be maintained at a proximal position to pivot the forefoot
segment, instead of the MTHs, for propulsive push-off (Brodtkorb et al., 2008).


114
4.3.2 Modification of insoles for foot-support interface applications
The FE model of a 6.8 mm thick, flat insole was specifically created and meshed
with eight-node hexahedron elements. An additional metatarsal support
component (Fig. 4.3.2.1) was merged into the flat insole model by projecting its
boundary nodes onto a curved surface that represents the geometrical shape of
an actual metatarsal support (Foot in Motion, USA).


Fig. 4.3.2.1 Finite element model of a flat insole with an integrated metatarsal support
component interfacing with the forefoot plantar surface

The analysis starts from the baseline case with the foot model landing on
a flat, rigid ground, and then increases complexity of interactions by adding a soft
insole integrated with an additional metatarsal support component. Parametrical
analysis was performed in order to determine the optimal material selection (soft
plasterzote versus stiffer EVA) and positioning (just underneath versus 5 mm or

12 mm proximal to the target MTH) of the metatarsal support for maximum MTH
unloading.



115
4.3.3 Efficacy of modified insole in relieving loads under MTHs

Fig. 4.3.3.1. Finite element predicted plantar anterio-posterior shear stresses with
different foot-supporting interfaces (A) rigid ground, (B) soft foam pad, (C) a soft insole
integrated with metatarsal support. Note that positive values denote anterior-directed AP
shear and medial-directed ML shear.

116
Based on the simulation results, landing on the rigid floor, the model predicted
peak plantar pressure and shear stress were 590.2 kPa and 205.9 kPa,
respectively. Simulations using the foam pad resulted in a regional peak pressure
reduction by 36.8% at the 2nd MTH (Table 4.3.3.1.). Peak resultant shear stress
found at this site was 109.8 kPa. An additional soft metatarsal support placed at
5 mm proximal to the 2nd MTH helped to reduce local peak pressure by 12.1%
as compared to using soft foam pad alone. However, placing a stiffer metatarsal
support just underneath the 2nd MTH could cause local peak pressure increase
to 504.4 kPa. Inclusion of a metatarsal support into the foam pad for enhanced
sub-MTH stresses relief requires more technical efforts.

Table 4.3.3.1. Model predicted changes in plantar stress peaks under 2nd MTH due to
different foot support conditions. (NA: no metatarsal support; MSP-5: metatarsal support
positioned 5 mm proximal to the 2nd MTH)
PPP (KPa) PSS-AP (KPa) PSS-ML (KPa)
Support

conditions
NA MSP-5 NA MSP-5 NA MSP-5
Rigid 590.2 503.2 17.9 82.1 55.7 63.3
EVA 472.7 466.6 38.2 27.3 23.6 2.2
plastazote 321.6 176.2 37.9 7.8 23.1 5.3

Plantar stress distributions were very sensitive to the metatarsal support’s
placement as well as its material selection. The current FE model provides an
efficient computational framework to investigate the efficacy of certain design
variables used in therapeutic footwear. Also, the model has the potential for
three-dimensional contact stress analysis at foot-shoe interface.