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The Immediate Effect of Foot Orthoses on Subtalar Joint Mechanics and Energetics

MAHARAJ, JAYISHNI N.; CRESSWELL, ANDREW G.; LICHTWARK, GLEN A.

Medicine & Science in Sports & Exercise: July 2018 - Volume 50 - Issue 7 - p 1449–1456
doi: 10.1249/MSS.0000000000001591
APPLIED SCIENCES
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Purpose Foot orthoses maybe used in the management of musculoskeletal disorders related to abnormal subtalar joint (STJ) pronation. However, the precise mechanical benefits of foot orthoses for preventing injuries associated with the STJ are not well understood. The aim of this study was to investigate the immediate effect of foot orthoses on the energy absorption requirements of the STJ and subsequently tibialis posterior (TP) muscle function.

Methods Eighteen asymptomatic subjects with a pes planus foot posture were prescribed custom-made foot orthoses made from a plaster cast impression. Participants walked at preferred and fast velocities barefoot, with athletic footwear and with athletic footwear plus orthoses, as three-dimensional motion capture, force data, and intramuscular electromyography of the TP muscle were simultaneously collected. Statistical parametric mapping was used to identify time periods across the stride cycle during which footwear with foot orthoses significantly differed to barefoot and footwear only.

Results During early stance, footwear alone and footwear with orthoses significantly reduced TP muscle activation (1%–12%), supination moments (3%–21%), and energy absorption (5%–12%) at the STJ, but had no effect on STJ pronation displacement.

Conclusions The changes in TP muscle activation and STJ energy absorption were primarily attributed to footwear because the addition of foot orthoses provided little additional effect. We speculate that these results are most likely a result of the compliant material properties of footwear. These results suggest that athletic footwear may be sufficient to absorb energy in the frontal plane and potentially reducing any benefit associated with the addition of foot orthoses.

School of Human Movement and Nutrition Sciences, The University of Queensland, Brisbane, Queensland, AUSTRALIA

Address for correspondence: Jayishni Maharaj, BHealthSc (Podiatry, Hons I), Ph.D., School of Human Movement and Nutrition Sciences, The University of Queensland, Brisbane, 4072 QLD, Australia; E-mail: jayishni.m@uq.edu.au.

Submitted for publication October 2017.

Accepted for publication February 2018.

Supplemental digital content is available for this article. Direct URL citations appear in the printed text and are provided in the HTML and PDF versions of this article on the journal’s Web site (www.acsm-msse.org).

Foot orthoses are commonly prescribed for the prevention and management of musculoskeletal disorders of the lower limb (1,2). The evidence base for the use of custom-made foot orthoses is incomplete (1); however, recent research suggests that orthoses may be effective in preventing musculoskeletal injuries (2) and beneficial for rear foot pain in rheumatoid arthritis (1,3). Foot orthoses are often advocated for musculoskeletal disorders because of their putative association with abnormal foot function, particularly in the frontal plane. The mechanistic effects of foot orthoses, however, are equivocal for resisting or facilitating motion (4,5), plantar pressure reduction (6,7), altered muscle activity (8–10), and enhanced postural stability (11) during walking and running.

During locomotion, the foot is required to absorb mechanical energy during the stance phase (12). This is achieved via compression of the longitudinal arch (sagittal plane [13]) and active resistance to subtalar joint (STJ) pronation (frontal plane [14]). The STJ absorbs energy via strain of the tibialis posterior (TP) tendinous tissue (i.e., tendon, aponeurosis and other connective tissue), while the TP muscle fascicles generate force more or less isometrically (14,15). The requirements of the TP tendinous tissue (referred to as TP tendon henceforth) to repetitively absorb energy during walking may predispose it to strain-induced tendinopathy. However, it is conceivable that foot orthoses, which are typically used in the management of TP tendon injuries (16), may reduce STJ moments and thus assist in the energy absorption requirements of the foot and reduce the strain of the TP tendon.

During walking, one potential function of foot orthoses is to reduce STJ moments (17) and act as a damper, between the foot and the ground (5). By applying a force medial to the STJ axis, foot orthoses can absorb energy and may reduce the forces generated by the TP muscle and subsequently the energy absorbed by the TP tendon. Orthoses have been shown to reduce STJ moments (18,19) and TP activation during early stance (9). However, no study to date has simultaneously measured STJ kinematics, kinetics, and TP muscle activation to understand the capacity of foot orthoses to absorb energy.

This study evaluated the immediate effects of custom-made foot orthoses, on frontal plane energy absorption during gait. We examined a homogeneous group with pes planus foot deformity using both barefoot and footwear walking as control conditions. We hypothesized that foot orthoses would reduce the energy absorbed in the frontal plane at the STJ by altering the forces required by the TP muscle (evident by a decrease in TP muscle activation and STJ supination moments). To determine the role of the STJ in absorbing energy and generating power during walking, intramuscular electromyography (EMG) of the TP to quantify muscle activation was simultaneously measured with multisegment foot kinematics. STJ mechanics were measured using a subject-specific multisegment foot model to quantify physiological triplanar motion of pronation and supination during walking.

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METHODS

Participants

Eighteen asymptomatic subjects (13 men, 5 women) with pes planus feet who had no injuries and not worn foot orthoses for a minimum of 2 yr gave written consent to participate in the study. The subjects’ average age, height, and mass were 26 ± 5 yr, 1.70 ± 0.11 m, 71.3 ± 12.6 kg, respectively. A priori calculations estimated a sample size of 15 participants to achieve a power of 80% (effect size = 0.8 [9], alpha level = 0.05) using a repeated-measures ANOVA analysis among three groups. The protocol was approved by the local university ethics committee and conducted according to the Declaration of Helsinki.

The foot posture index (FPI) was used to categorize foot posture. To qualify, participants had to exhibit a total score of +6 or greater (20). Each participant’s FPI was measured in a relaxed standing position (by investigator J.N.M.) and was scored using six criteria on a 5-point scale from −2 to +2 (20). The six scores were summed to give each subject a composite score between −12 and +12. The FPI has been shown to have good intrarater reliability (21).

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Foot orthoses

A plaster cast impression was taken of each participant’s feet in the STJ neutral position using the suspension technique by a podiatrist (investigator J.N.M.) with 5 yr of experience (22). A commercial orthotic laboratory (Queensland Orthotic Laboratory, Brisbane, Australia) custom-manufactured pairs of three-quarter length foot orthoses with a 4-mm medial skive at 15° and a 5° extrinsic rear foot post (23,24) (Fig. 1). The orthoses were made from a semirigid 4-mm polypropylene thermoplastic shell with vinyl covering, characteristics of a commonly prescribed device to reduce STJ moments in Australia (25). All patients were given a minimum 15-min period of acclimatization with the orthoses in a standardized shoe before data collection.

FIGURE 1

FIGURE 1

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Experimental protocol

Participants walked on a force-instrumented tandem treadmill (DBCEEWI; AMTI, Watertown, MA), while kinematic, kinetic, and EMG data were synchronously recorded in three conditions (barefoot, athletic footwear, and athletic footwear with orthoses (the latter condition being subsequently referred to as “orthoses”). A standardized athletic shoe (Asics Gel Lyte 33; Asics Corp., Tokyo, Japan) was used for all subjects. Participants selected their preferred walking velocity as the speed of the treadmill was incrementally increased while wearing both footwear and orthoses. This was repeated as the speed was decreased and until the subject consistently identified the same preferred velocity three times. Participants were tested at preferred and fast (40% faster than preferred) walking velocities in each condition, in a randomized order. At each speed, participants were given 1 min to normalize their gait before a 10-s period of data collection. The average group velocity during preferred and fast walking was 1.2 and 1.7 m·s−1, respectively.

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Joint mechanics

The procedures for collecting and processing STJ kinematics and kinetics have been reported earlier by Maharaj et al. (26). Briefly, STJ kinematics and kinetics were calculated using subject-specific models scaled from a custom built multisegment foot model in OpenSim software (27). The STJ of the three-dimensional musculoskeletal model (26,28) was modeled as a revolute joint with an axis of 38° inclination and 21° deviation to the midline of the body, running from posterior, inferior and lateral to anterior, superior and medial through the rear foot (29). The modeled oblique axis allows the STJ to perform triplanar motion of pronation and supination. The calculated STJ moments represent the net internal moments produced by the musculoskeletal structures at the STJ. STJ displacement was calculated relative to the angle at heel strike, and STJ power was calculated as the product of STJ moment and STJ angular velocity. To calculate STJ negative and positive work, STJ power was integrated with respect to time using trapezoidal integration. Moments, velocities, power, and work measures were normalized to nondimensional forms using foot length (m) and body weight (N) (30,31).

Step length was defined as the anterior–posterior distance in the direction of progression between the right and left distal calcaneal markers. Step width was defined as the mediolateral distance between the right and left distal calcaneal markers. Step length and width were normalized to the length of the leg (distance between the superior iliac spine and medial malleolus markers).

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Muscle activation

TP muscle activation was collected using intramuscular EMG for which the procedures are reported by Maharaj et al. (14). In short, two fine-wire electrodes (0.051-mm stainless steel, Teflon- coated; Chalgren Enterprises, Gilroy, CA) were inserted into the muscle in a bipolar configuration using a posterior–medial approach under ultrasound guidance (Echoblaster, 128, UAB; Telemed, Vilnius, Lithuania). After DC-offset removal, the TP signal was high-pass filtered using a fourth-order Butterworth filter at 30 Hz. To generate an EMG envelope, an RMS amplitude was calculated using a moving window of 100 ms. Subsequently, the EMG envelope was normalized to its maximum RMS amplitude during the preferred walking condition. Only EMG data from 14 of the 18 participants were analyzed owing to technical difficulties in data collection.

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Statistics

Statistical parametric mapping (32) was used to assess differences between the barefoot, footwear, and orthoses conditions for STJ displacement, normalized STJ moments, normalized STJ power, and normalized TP EMG. In summary, the method uses random field theory to make statistical inferences about continuous one-dimensional time series data to test where signals may differ in the time domain. Hypothesis testing in a continuous manner is valuable because it takes into account the dependency between different time instances of the gait cycle and reduces post hoc regional focus bias (33). A mean of the time normalized kinematic, kinetic, and EMG data from a minimum of three strides was calculated for each subject in each condition. Statistical parametric map (SPM) one-dimensional package in Matlab (available at http://www.spm1d.org) was used to perform a two-way repeated-measures ANOVA to describe the main effects of support condition (barefoot, footwear, orthoses) and velocity (preferred and fast) on STJ displacement, moments, power, and TP EMG, as well as any interaction effects (support condition–velocity).

For each outcome variable, three statistical parametric maps (SPM{F}) were created for each independent factor and their interactions (support condition, velocity and interactions) by calculating the conventional F statistic at each time point of the stride cycle (see Figure, Supplemental Digital Content 1, SPM repeated-measures two-way ANOVA results for STJ displacement (A), normalized STJ moments (B), normalized STJ powers (C), and normalized TP EMG (D), http://links.lww.com/MSS/B219). If the SPM{F} crossed the critical threshold estimated using random field theory (34), it would indicate a significant effect of the independent factors or an interaction. When significant main or interaction effects were found, post hoc paired t-tests including Bonferroni corrections (33) were performed with the footwear condition (i.e., barefoot vs footwear and footwear vs orthoses) at each velocity. If at any time, SPM{t} crossed the critical threshold, a suprathreshold cluster was created, indicating a significant difference between support conditions at the specific location of the gait cycle.

SPM zero-dimensional package was used to perform discrete parameters two-way repeated-measures ANOVA to describe the effects of support condition, velocity, and interactions (support condition–velocity) on normalized net STJ work, normalized step length, and normalized step width. Statistical significance was established at P ≤ 0.05. Effect size (h2) was computed as a ratio of the effect variance (SSeffect) to the total variance (SStotal). Periods of significant difference between conditions are indicated by black bar at the top of each graph. Results are presented as mean ± SD unless otherwise stated.

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RESULTS

At the preferred velocity, there was a significant main effect of support condition on step length (Fig. 2A; ANOVA, effect of support condition; F2,26 = 49.53; P ≤ 0.01, h2 = 0.6295), which was not evident at the fast velocity (significant support condition–velocity interaction, P ≤ 0.01). Post hoc analysis revealed that footwear significantly reduced the step length compared with barefoot walking (adjusted P ≤ 0.01). Step width remained consistent across the three conditions (Fig. 2B; ANOVA, effect of support condition; F2,26 = 1.455; P ≥ 0.05). There was a significant main effect of velocity on step width (ANOVA, effect of velocity; F1,13 = 6.238; P ≤ 0.05, h2 = 0.0823), such that step width decreased during fast walking (Fig. 2B).

FIGURE 2

FIGURE 2

The net work generated by the STJ (Fig. 2C) was significantly different across the three conditions (ANOVA, effect of support condition; F2,34 = 4.57; P ≤ 0.01, h2 = 0.1740) and across speeds (ANOVA, effect of support condition; F1,17 = 20.84; P ≤ 0.01, h2 = 0.3996). The greatest reduction in net work relative to the barefoot condition was due to the effects of footwear alone (adjusted P ≤ 0.05).

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Effect of support condition

In early stance, there was a significant main effect of support condition on TP muscle activity between 1% and 12% (P ≤ 0.01), STJ moments between 3% and 21% (P ≤ 0.01), and STJ power between 5% and 12%, (P ≤ 0.01) of the stride cycle, such that there was reduced muscle activation, reduced STJ supination moment, and reduced STJ negative power absorbed in the footwear and orthoses conditions (Fig. 3A–D). Later in stance, a significant decrease in TP muscle activity during 19%–22% and 32%–44% (P ≤ 0.01), STJ angular displacement during 45%–59% (P ≤ 0.01), STJ supination moment during 55%–64% (P ≤ 0.01), and STJ positive power during 24%–28%, 37%–44%, and 47%–51% (P ≤ 0.01) of the stride cycle was observed in the footwear and orthoses conditions (Fig. 3A–D).

FIGURE 3

FIGURE 3

Multiple comparison analyses highlighted that the effects of support condition were primarily a result of footwear, with no difference between footwear and orthoses conditions. During early stance, footwear significantly reduced TP muscle activity between 8% and 11% (P ≤ 0.01; Fig. 3D) and the power absorbed between 7% and 11% (P ≤ 0.01; Fig. 3C) of the stride cycle compared with the barefoot condition. Later in stance, footwear significantly reduced STJ moments over 32%–45% and 56%–63% (P ≤ 0.01) of the stride cycle; however, no main effect of support condition was highlighted during the period.

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Effect of velocity

In early stance, an increase in walking velocity resulted in a significant increase in supination displacement over 1%–6% (P ≤ 0.01), TP muscle activity over 1%–4% and 6%–11% (P ≤ 0.01), STJ supination moment over 8%–17% (P ≤ 0.01), and STJ power over 6%–12% (P ≤ 0.01) of the stride cycle. During the second half of stance, an increase in TP muscle activity between 31% and 44% (P ≤ 0.01), STJ supination moment between 54% and 63% (P ≤ 0.01), STJ angular displacement between 46% and 61% (P ≤ 0.01), and STJ power generation between 23% and 30% and 42% and 52% (P ≤ 0.01) of the stride cycle was evident as a result of an increase in velocity (Fig. 3E–H).

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Support condition and velocity interactions

There were significant interaction effects (support condition–velocity) on TP muscle activity over 5%–9% (P = 0.02) and supination moment over 3%–18% (P ≤ 0.01) of the stride cycle. These results suggest that during these periods in early stance, the effect of support condition on TP muscle activity and STJ moments only occurred at the faster velocity (Fig. 4).

FIGURE 4

FIGURE 4

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DISCUSSION

The objective of this study was to evaluate the immediate effects of footwear plus custom-made foot orthoses in absorbing frontal plane mechanical energy during walking in a healthy population with a pes planus foot structure. Footwear and footwear plus orthoses reduced the power absorbed at the STJ in early stance. This was achieved through a reduction in TP muscle activation and consequently the supination moment generated about the STJ joint during weight acceptance. A significant reduction in power generation and TP muscle activation was also evident during late stance when wearing footwear with and without orthoses. Walking in footwear and orthoses at faster velocities demonstrated similar power profiles at the STJ, although greater magnitudes of power absorption and generation were required compared with the slower preferred walking velocity. Of particular interest was the lack of significant mechanical and energetic differences between the footwear and orthoses conditions at both velocities, which we speculate is a result of the material properties of the orthotic device.

The viscoelastic materials in athletic footwear, including ethylene-vinyl acetate (EVA), polyurethane, and various forms of gel capsules, are designed to attenuate shock and absorb energy. Compliant shoes may absorb energy by reducing peak forces and decreasing peak plantar pressures by distributing load over a greater area (35). The sock liner of the shoe, which is typically manufactured from EVA, may also provide shock attenuation. In healthy participants, flat EVA insoles have been shown to significantly reduce peak plantar pressures at the heel to a similar degree to custom-made foot orthoses (36). Although stiffness properties of the footwear and orthoses were not quantified in this study, we believe that the semirigid 4-mm polypropylene orthoses were significantly stiffer than the footwear and thus failed to provide additional energy absorption during walking.

In early stance, footwear likely reduced the energy absorbed in the TP tendon by decreasing the force generated by the muscle (concurrent decrease in TP activation and peak STJ supination moment), while the TP tendon is known to stretch to absorb the energy at the STJ (14). If less energy is stored in the TP tendon, then this would subsequently diminish the ability of TP to generate power during late stance, which is supported by our results. Tendons are efficient biological springs, returning approximately 90% to 97% of the energy that they store (37), while footwear materials lose approximately 30% to 40% of their absorbed energy (38). Nonbiological materials also lack the ability to transfer energy. For example, plastics and EVA cannot reuse the energy stored via deformation at the heel to generate power at the forefoot in late stance. Energy is lost in footwear and must therefore be compensated by active force generation. These forces are mostly likely generated by the plantar flexors and the intrinsic foot muscles (39), although it may not be as energetically efficient as elastic power generation.

Spatiotemporal measures of walking may also contribute to the energy absorption and generation requirements of the STJ with footwear and orthoses. At preferred velocity, significantly shorter step lengths were evident with footwear and orthoses compared with barefoot walking. Shorter steps require less mechanical work to redirect the center-of-mass velocity from the trailing limb to the leading limb and thus are beneficial for the mechanical and metabolic efficiency of gait (40). The shorter step lengths in footwear compared with barefoot are contrary to previous research, which find that step length is typically shorter in barefoot conditions when walking velocity is self-selected (41). The discrepancy is likely to be a result of the protocol used here. Specifically, in our protocol, preferred walking velocity was determined in the orthoses condition, which is likely to be greater than the preferred velocity when walking barefoot (42). Thus, to accommodate the higher velocity, a greater step length and frequency may have been required in the barefoot condition. Nevertheless, at the faster walking velocity, the differences in step length disappeared, yet the significant reductions in energy absorption and generation, moments, and activations prevailed. In fact, the interaction effect suggests that reductions in the joint moments and muscle activations were primarily evident at the faster velocity, highlighting the significant effect of footwear.

At the fast walking velocity, the greater magnitudes of power absorption and generation across the three conditions compared with the slower preferred walking velocity are most likely a result of a reduced step width. With a narrower step width, a significant increase in TP muscle activation and STJ moment was evident through stance. The narrow step width is likely to increase the external moment arm of the STJ relative to the ground reaction force during walking, and thus, the moments generated. Although the effect of step width on STJ moment arm is speculatively, the association between a narrower step width and greater STJ moments during gait has been previously demonstrated (26). A narrow step width has also been shown to reduce mediolateral stability during walking (43), suggesting that greater active force generation by the TP muscle (evident by the greater TP muscle activation, STJ moment) at the fast walking velocity may be required to maintain mediolateral stability. These results highlight the critical role of the TP in maintaining frontal plane mechanics of the foot during walking. The greater energy absorption requirements at faster walking velocities also suggest greater strain and risk of injury in the TP tendon (15).

Athletic footwear with compliant material properties may be beneficial in the management of pathologies associated with pes planus deformity (such as TP tendinopathy), particular when walking at faster velocities because of their capacity to absorb power and reduce the net work performed at the STJ during walking. The associated decrease in TP stretch with footwear and foot orthoses may prevent further microtrauma in a painful TP tendon and therefore facilitate greater load tolerance (16). This could allow individuals to maintain daily activities such as standing and walking. Our results also suggest that altering step width during gait may aid in reducing the STJ supination moments. These speculations warrant further studies evaluating the longer-term efficacy of foot orthoses for TP tendinopathy and the biomechanical and energy dissipation mechanisms of foot orthoses on the STJ joint.

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Methodological considerations

There are some limitations and assumptions of the present methodology that should be acknowledged. First, the acclimatization time of 15 min may not have been sufficient to enable the neuromotor system to respond optimally to foot orthoses, preventing the foot orthoses from further altering muscle activity compared with footwear only. Because participants habitually walk barefoot or with footwear, they are more likely to be acclimatized to these conditions than the footwear plus orthoses condition. Murley et al. (9) found significant differences in TP muscle activation between footwear and foot orthoses conditions with an acclimatization period of approximately 7 days. Second, FPI may not be the most suitable measure to predict STJ mechanics during walking (26). Although there is evidence to suggest that FPI is associated with kinematic foot function (44), very recent research from our laboratory demonstrates that FPI is not a significant predictor of STJ kinetics and energy absorption during walking (26). Furthermore, clinicians may criticize the standardized nature of the foot orthoses; although we believe that the prescription of the foot orthoses is well justified. All subjects included in the study were categorized with a planus foot deformity, and orthoses were designed to maximize rear foot support. In addition, although the speeds at each condition were randomized, the support conditions were not; thus, an order effect may be present. However, we are confident that the primarily effects of order on fatigue and performance were minimized by allowing participants to rest between each support condition and the initially 15-min acclimatization period on the treadmill with footwear and orthoses.

To calculate moments about the STJ, we performed inverse dynamics using a custom-built foot model with a fixed axis of rotation about the STJ (26). Although the STJ axis has been previously shown not to be fixed (45), previous research from our laboratory suggests that the physiological deviations in the STJ axis have minimal effect on the negative work absorbed at the STJ (26). However, because changes to the inclination angle may moderately influence the moments generated about the STJ, the ankle joint was given multiple degrees of freedom, allowing the STJ axis to move relative to the tibia, in a manner similar to a moving STJ axis. It should also be acknowledged that STJ moments are likely to be contributed to by muscles other than the TP; for example, the triceps surae, tibialis anterior, and peroneus longus. As such, changes in STJ moment may also be the result of changes in forces in these muscles. However, because TP has the greatest supination moment arm at the STJ (46) and a relatively large physiological cross-sectional area, it is likely to have the greatest influence at the STJ. Therefore, any changes in STJ mechanical function with footwear and/or orthoses are likely to primarily affect TP muscle.

The authors would like to acknowledge Australian Podiatry Education and Research Foundation for their financial contribution and ASICS Oceania for providing footwear in kind. J. N. Maharaj was supported by a National Health and Medical Research Council postgraduate scholarship (APP1075000). The authors declare no conflict of interest.

The authors declare that the results of the study are presented clearly, honestly, and without fabrication, falsification, or inappropriate data manipulation. The results of the present study do not constitute endorsement by the American College of Sports Medicine.

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Keywords:

FOOT ORTHOTICS; FOOTWEAR; FRONTAL PLANE; WALKING; ENERGETICS

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