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The influence of great toe valgus on pronation and frontal plane knee motion during running

Posted on March 31, 2020 | Comments Off on The influence of great toe valgus on pronation and frontal plane knee motion during running

by Richard Stoneham PhD1, Gillian Barry PhD1, Lee Saxby BSc2, Mick Wilkinson PhD1*

The Foot and Ankle Online Journal 13 (1): 7

Injury rates in running range from 19.4‐79.3%, with injuries at the knee comprising 42.1%. Pronation and altered frontal plane knee joint range of motion have been linked to such injuries. The influence of foot structure on pronation and knee kinematics has not been examined in running. This study examined associations between great toe valgus angle, peak pronation angle and frontal plane range of movement at the knee joint during overground running while barefoot. Great toe valgus angle while standing, and peak pronation angle and frontal plane range of motion of the dominant leg during stance while running barefoot on an indoor track were recorded in fifteen recreational runners. There was a large, negative association between great toe valgus angle and peak pronation angle (r = -0.52, p = 0.04), and a strong positive association between great toe valgus angle and frontal plane range of motion at the knee joint (r = 0.67, p = 0.006). The results suggest that great toe position plays an important role in foot stability and upstream knee-joint motion. The role of forefoot structure as a factor for knee-joint injury has received little attention and could be a fruitful line of enquiry in the exploration of factors underpinning running-related knee injuries.

Keywords: great toe valgus, pronation, frontal plane knee range of motion, running

ISSN 1941-6806
doi: 10.3827/faoj.2018.1301.0007

1 – Department of Sport, Exercise and Rehabilitation, Northumbria University, UK
2 – LeeSaxby.com, Suffolk House, Louth, Lincolnshire, UK
* – Corresponding author: mic.wilkinson@northumbria.ac.uk

Injury incidence in running ranges from 19.4‐79.3% [1, 2]. The knee is the most injured site, comprising 42.1% of all running‐related injuries [2, 3]. Patellofemoral Pain Syndrome (PFPS) is the most common running‐related knee injury, followed closely by Iliotibial Band Syndrome (ITBS) [3].  Altered frontal plane hip and knee joint kinematics and pronation during the stance phase of running have been linked to these injury types, and differentiate injured from uninjured runners [4-6]. Knee abduction, femoral internal rotation, tibial external rotation, and foot pronation, have been theoretically linked to injury in a population of patients with PFPS [7]. As such, interventions to normalise altered frontal plane kinematics during running might be valuable for rehabilitation of this type of knee injury. Interventions have tended to focus on proximal areas linked to altered knee kinematics. However, training studies to increase hip abduction and external rotation strength have not decreased hip or knee frontal plane peak joint angles or joint excursions during the stance phase of running [8-10]. Moreover, associations between hip strength and frontal plane hip and knee peak angles and joint excursions while running and jumping are weak [9, 11]. These findings suggest that proximally-based interventions are not effective at altering lower extremity running mechanics and risk of running related injury. Studies exploring the distal end of the kinetic chain have utilised barefoot and minimal footwear, and hip and foot muscle strengthening interventions to reduce surrogate measures associated with injury at the knee and other sites [10, 12-14]. Injury rates, however, remain high [15]. The influence of foot structure on pronation and knee joint kinematics in running has, by contrast, received little attention.

Data comparing foot structure in habitually-barefoot and habitually-shod populations have reported consistent differences in the spread/abduction of the great toe from the other toes [16-19]. Based on Newtonian physics, larger areas of support provide greater stability. It has been suggested that an abducted great toe position might be important for controlling the direction of body weight during running, secondary to improved stability of the foot [20, 21]. Running is essentially a series of alternate single-leg jumps, where multiples of bodyweight must be supported and controlled using a spring-like action of the supporting foot and limb [22, 23]. Early research showed an active role of the toes, the great toe in particular, from midstance to toe off in running [24]. More recent data comparing habitually barefoot to habitually shod populations suggested that the abducted great toe position, characteristic of the barefoot group, reduced peak forefoot pressures during running by increasing the area of support [19]. Another comparative study from the same lab [25] found larger ankle eversion and internal rotation (which together comprise pronation) during the landing phase of jumping in habitually shod compared to habitually-barefoot participants, attributing differences to the abducted great toe position characteristic of the barefoot group. Together, these studies suggest a link between great toe position and foot and ankle stability in running, and dynamic tasks with similar demands to running. Given evidence of the link between pronation, altered frontal plane motion at the knee joint and risk of knee injury [7], there is a possible mechanistic link between great toe position, pronation and frontal plane knee joint kinematics.

Previous research suggests that the toes have a stabilising function, and that great toe position influences area of support in running, and the extent of pronation in the landing phase of jumping. The influence of great toe position on pronation and on kinematics at the knee joint has not been examined in running. The aim of this study was to examine associations between great toe valgus angle, peak pronation angle and frontal plane range of movement at the knee joint during overground running while barefoot, the latter being necessary to avoid toe position being constrained by shoes.

Methods

Participants

With institutional ethics approval, 15 volunteers (ten male, five female) participated. Mean and SD age, stature and mass of all participants were 26±7 yrs, 1.71±0.01 m and 69±10.9 kg respectively. Inclusion criteria were aged 18-45 years and participation in endurance running more than once per week as part of habitual-exercise regimes, with at least one run longer than 30 minutes. Participants were excluded if they had an injury to the lower limbs in the previous six months, or any condition that could affect their normal running gait.

Design

An observational design assessed the relationship between great toe valgus angle relative to the first metatarsal, peak pronation angle and frontal plane range of movement at the knee joint of the dominant leg during stance, while running barefoot on an indoor runway. The barefoot condition was chosen as it was the only way to ensure that the toe angle recorded in standing was not altered by footwear while running. Data were collected in a single visit. Participants were provided with a short-sleeved compression top and shorts to improve skeletal representation in biomechanical modelling, and were instructed to be well rested before testing. Reflective markers were attached in ‘Plug-In gait’ and ‘Oxford-Foot Model’ formations to assess lower-limb kinematics of the dominant limb. Participants were habituated to running barefoot with a 30-minute, self-paced run. After habituation, participants ran over a 20-m runway where kinematic data were captured by 14 optoelectronic cameras. Electronic timing gates (Brower timing gates, Utah, USA) placed in the data capture area (2.7m apart) were used to record speed in each trial. The average running speed was 2.48±0.38 m·s-1.

Procedures

Great toe valgus angle

Participants stood barefoot on top of a 0.35-m high platform covered in graph paper. The non-dominant foot was placed on the platform first, aligning the most posterior aspect with a horizontal reference line on the graph paper. The dominant foot was positioned next, shoulder width apart from the other foot, and with the most posterior aspect on the same horizontal reference line. The first metatarsal proximal-and distal-dorsal protrusions, and the central and dorsal point of the interphalangeal joint of the great toe were identified by palpation, and marked using a permanent pen. A digital camera (CX240, Sony, Japan) positioned 0.3m above the platform on a tripod was aligned with the first metatarsophalangeal joint, and the zoom was adjusted so that bony prominences defining great toe angle were visible. A still image was captured and saved for analysis of great toe valgus angle.

Kinematics

Prior to habituation, anthropometric measures were recorded for use in biomechanical modelling (stature (mm), mass (kg), bilateral-leg length (mm), and knee and ankle joint width (mm)). For assessment of lower-limb joint kinematics, participants had a series of markers (Ø=14mm) attached in ‘Plug-In gait’ and ‘Oxford-Foot Model’ formations. Anatomical locations of the ‘Plug-In gait’ and ‘Oxford-Foot Model’ were sacrum, bilateral anterior-and posterior-superior iliac spines, the bilateral distal-lateral thigh, bilateral femoral-lateral epicondyle, the bilateral distal-lateral lower-leg, the bilateral lateral malleoli, the left/right toe (dorsal aspect of the second metatarsal head) and the calcaneus of the non-dominant limb at the same height as the toe marker. The following markers were placed on the dominant limb only, lateral head of the fibula, tibial tuberosity, anterior aspect of the shin, the medial malleoli, the proximal aspect of the calcaneus, a ‘peg marker’ extending from the most posterior aspect of the calcaneus, the inferior aspect of the calcaneus, sustentaculum tali, proximal and dorsal aspect of the first metatarsal head, the medial and distal aspect of the first metatarsal head, the proximal-and distal-lateral aspects of the fifth metatarsal and the medial aspect of the first phalanx. Fourteen infrared-optoelectronic cameras (Vicon 10 xT20 and 2 x T40, Oxford, UK) captured kinematic trajectories at 200Hz. 

Data treatment

A trial was deemed successful when running speed was ± 5% of the predetermined running speed from the habituation run. Dominant limb data for peak pronation angle and frontal plane range of motion at the knee joint were exported to Microsoft Excel (Microsoft, USA). Foot structure images were loaded to Dartfish ClassroomPlus (version 7.0, Fribourg, Switzerland) where great toe valgus angle was measured using the angle tool. (Chicago, USA).

Statistical analysis

Statistical analysis was undertaken using JASP 0.10.2. Following verification of assumptions of linearity and uniformity of errors using Q-Q and residuals versus predicted value plots respectively, linear regression assessed associations between great toe valgus angle, peak pronation angle and frontal plane range of motion at the knee joint. Strength of associations were judged against Cohen’s effect size categories for Pearson’s  r i.e. small association 0.1-0.3; moderate association 0.3-0.5; large association 0.5-1.0 [26]   Significance was accepted at p < 0.05.

Results

Mean and SD great toe valgus angle, peak pronation angle and frontal plane knee range of motion were 9.5±6.1°, -5.2±6.6° and 6.2±2.2° respectively.

Association between great toe valgus and peak pronation angle.

There was a large, negative association of great toe valgus angle and peak pronation angle during stance (r = -0.52, p = 0.04). As great toe valgus angle increased (more positive = more valgus), peak pronation angle decreased (more negative = increased pronation) (see Figure 1). The regression equation showed a 0.59° increase in peak pronation for every additional degree of great toe valgus (95% CI 0.01 to 1.12°).

Figure 1 Association between great toe valgus angle and peak pronation angle during overground barefoot running on an indoor track in 15 recreational runners.

Association between great toe valgus and frontal plane knee range of motion.

Great toe valgus angle was strongly and positively associated with frontal plane range of motion at the knee joint (r = 0.67, p = 0.006). As great toe valgus angle increased, frontal plane knee range of motion also increased (see Figure 2). The regression equation showed a 0.24° increase in frontal plane knee joint excursion for every one degree increase in great toe valgus angle (95% CI 0.01 to 0.40°).

Figure 2 Association between great toe valgus angle and frontal plane range of motion at the knee joint during overground barefoot running on an indoor track in 15 recreational runners.

Discussion

The aim of this study was to examine associations between great toe valgus, peak pronation and frontal plane range of motion at the knee joint during overground running. There was a strong, negative correlation between great toe valgus angle and peak pronation such that increased great toe valgus was associated with a more negative peak pronation angle (increased pronation). There was also a strong, positive correlation between great toe valgus angle and frontal plane range of motion at the knee joint such that increased great toe valgus was associated with larger knee joint excursions in the frontal plane. Altered frontal plane hip and knee joint kinematics and pronation during the stance phase of running have been linked to running-related knee injury, and can differentiate injured from uninjured runners [4-6]. Knee abduction and foot pronation have also been theoretically linked to patellofemoral pain [7]. In light of this evidence, our results suggest that forefoot structure might be an important but largely unexplored factor in running-related knee injury.

As this is the first study to explore the association between great toe valgus, pronation and frontal plane knee joint excursions during running, there are no studies with a similar approach for comparison. Nevertheless, the strong relationships observed broadly support findings from previous comparative cross-sectional studies of habitually barefoot and habitually shod participants that differed in forefoot structure with respect to the spread/abduction of the great toe [19, 25]. Shu et al. [25] observed larger ankle eversion and internal rotation (which together comprise pronation) in habitually shod compared to habitually barefoot participants in the landing phase of jumping. As running is essentially a series of single-leg jumps, the strong association of great toe valgus angle with peak pronation observed in running in our study is not surprising. The reduced and more evenly distributed forefoot peak pressures of habitually barefoot participants reported by Mei et al. [19] alludes to greater forefoot stability during the period of stance when forces are highest. It is possible that as the stability provided by the great toe decreases with increasing valgus angle, instability of the foot could manifest as higher peak pronation. Increased forefoot instability with increased great toe valgus is a plausible mechanism that could explain the strong correlation of great toe valgus angle and peak pronation that we observed. Increased postural instability with great toe valgus [27] and with splinting of the great toe into flexion [28] have been observed in single-leg balance tasks. Though these studies examined static balance and not the dynamic single-leg balance characteristic of running, the underpinning link between the area of the base of support and subsequent stability could be assumed to be common to both. Instability at the foot could have kinematic consequences further up the kinetic chain, resulting in increased frontal plane motion at the knee. The strong, positive association of great toe valgus angle with frontal plane knee joint excursion observed in the current study is consistent with this suggestion. Moreover, the kinematic links between pronation and frontal plane knee joint range, as well as the link between these factors and running-related knee injury suggested here have been suggested previously elsewhere [7] and supported by previous studies [4-6].

The main limitation of this study is that the correlational design prevents any suggestion of a causal link between great toe valgus, peak pronation and frontal plane knee joint excursions. Another limitation is that great toe valgus angle was measured during static stance, not while running, so an assumption that valgus angle remains relatively unchanged when the foot is loaded during running is implicit in the interpretation of the results. Previous research, however, suffers from similar limitations, comprising only comparative studies of foot and ankle function and pressure distributions of groups with mean abducted versus mean valgus great toe positions. As such, a correlational study like this one does add to the understanding of how foot structure might relate to pronation and knee joint kinematics in dynamic tasks like running by examining a ‘dose-response’ type association, in addition to the ‘with and without’ type evidence of previous comparative studies. Moreover, there are plausible mechanisms of action for both key findings in this study, so the data provide both direct and mechanistic evidence towards establishing a causal link [29]. A logical next step for this area of research would be randomised control trials where pronation and knee kinematics are evaluated before and after an intervention to alter great toe valgus angle in one group, with the control group foot structure remaining unchanged. Interventions could potentially include corrective surgery or corrective devices that reposition the great toe. Additional comparative studies that measure knee joint kinematics during running would, however, be a useful intermediate step.

In summary, this study observed strong associations between great toe position, peak pronation and frontal plane range of motion at the knee joint during over-ground barefoot running. The results suggest that great toe position plays an important role in foot stability and subsequent knee-joint motion. Both pronation and frontal plane knee-joint motion have been implicated in the etiology of knee injuries. The role of forefoot structure as a factor for knee-joint injury has received little attention and could be a fruitful line of enquiry in the exploration of factors underpinning running-related injuries.

This study formed part of a PhD program collaboratively funded by Northumbria University and VivoBarefoot. VivoBarefoot had no input to the design, analysis or interpretation of studies or data, or the preparation of this manuscript.

References

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Effects of medial and lateral orthoses on kinetics and tibiocalcaneal kinematics in male runners

Posted on December 31, 2017 | Comments Off on Effects of medial and lateral orthoses on kinetics and tibiocalcaneal kinematics in male runners

by Jonathan Sinclair1*

The Foot and Ankle Online Journal 10 (4): 1

Background: The aim of the current investigation was to examine the effects of foot orthotic devices with a 5° medial and lateral wedge on kinetics and tibiocalcaneal kinematics during the stance phase of running.
Material and methods: Twelve male participants ran over a force platform at 4.0 m/s in three different conditions (5° medial orthotic, 5° lateral orthotic and no-orthotic). Tibiocalcaneal kinematics were collected using an 8 camera motion capture system and axial tibial accelerations were obtained via an accelerometer mounted to the distal tibia. Biomechanical differences between orthotic conditions were examined using one-way repeated measures of analysis of variance (ANOVA).
Results: The results showed that no differences (P>0.05) in kinetics/tibial accelerations were evident between orthotic conditions. However, it was revealed that the medial orthotic significantly (P<0.05) reduced peak ankle eversion and relative tibial internal rotation range of motion (-10.75 & 4.98°) in relation to the lateral (-14.11 & 6.14°) and no-orthotic (-12.37 & 7.47°) conditions.
Conclusions: The findings from this study indicate, therefore, that medial orthoses may be effective in attenuating tibiocalcaneal kinematic risk factors linked to the etiology of chronic pathologies in runners.

Keywords: running, biomechanics, orthoses, kinetics, kinematics

ISSN 1941-6806
doi: 10.3827/faoj.2017.1004.0001

1 – Center for Applied Sport Exercise and Nutritional Sciences, School of Sport and Wellbeing, Faculty of Health & Wellbeing, University of Central Lancashire, Preston, Lancashire, PR1 2HE.
* – Corresponding author: jksinclair@uclan.ac.uk

Distance running is associated with a significant number of physiological and psychological benefits [1]. However, epidemiological analyses have demonstrated that pathologies of a chronic nature are extremely common in both recreational and competitive runners [2] and as many as 80% of runners will experience a chronic injury as a consequence of their training over a one-year period [2].

Given the high incidence of chronic pathologies in runners, a range of strategies have been investigated and implemented in clinical research in an attempt to mitigate the risk of injury in runners. Foot orthoses are very popular devices that are used extensively by runners [3]. It has been proposed that foot orthoses may be able to attenuate the parameters linked to the etiology of injury in runners, thus they have been cited as a mechanism by which injuries can be prophylactically avoided and also retrospectively treated [4]. The majority of research investigating the biomechanical effects of foot orthoses during running has examined either impact loading or rearfoot eversion parameters which have been linked to the etiology of running injuries. Sinclair et al, [5] showed that an off the shelf orthotic device significantly reduced vertical rates of loading and axial tibial accelerations, but did not alter the magnitude of rearfoot eversion. Butler et al, [6] examined three-dimensional (3D) kinematic/ kinetic data alongside axial tibial accelerations during running, using dual-purpose and a rigid orthoses. Their findings revealed that none of the experimental parameters were differed significantly between the different orthotic conditions.  Laughton et al, [7] showed that foot orthoses significantly reduced the loading rate of the vertical ground reaction force but did not significantly influence rearfoot eversion parameters. Dixon, [8] examined the influence of off the shelf foot orthoses placed inside an military boot on kinetic and 3D kinematic parameters during running. The findings from this investigation revealed that the orthotic device significantly reduced the vertical rate of loading, but no alterations in ankle eversion were reported.

Further to this, because the mechanics of the foot alter the kinetics/kinematics of the proximal lower extremity joints, biomechanical control of the foot with in-shoe orthotic wedges has wide-ranging applications for the treatment of a variety chronic lower extremity conditions. Different combinations of wedges or posts have therefore been used in clinical practice/ research to treat a multitude of chronic pathologies [9]. Both valgus (lateral) and varus (medial) orthoses have been proposed as potentially important low-cost devices for the conservative management of chronic pathologies [10].

Lateral orthoses are utilized extensively in order to reduce the loads experienced by the medial tibiofemoral compartment [10]. Lateral orthoses cause the center of pressure to shift medially thereby moving the medial-lateral ground reaction force vector closer to the knee joint center [11]. This serves to reduce the magnitude of the knee adduction moment which is indicative of compressive loading of the medial aspect of the tibiofemoral joint and its progressive degeneration [12]. Kakihana et al, investigated the biomechanical effects of lateral wedge orthoses on knee joint moments during gait in elderly participants with and without knee osteoarthritis [13]. The lateral wedge significantly reduced the knee adduction moment in both groups when compared with no wedge. Butler et al, examined the effects of a laterally wedged foot orthosis on knee mechanics in patients with medial knee osteoarthritis [14]. The laterally wedged orthotic device significantly reduced the peak adduction moment and also the knee adduction excursion from heel strike to peak adduction compared to the non-wedged device. Kakihana et al, examined the kinematic and kinetic effects of a lateral wedge insole on knee joint mechanics during gait in healthy adults [15]. The wedged insole significantly reduced the knee adduction moment during gait in comparison to the no-wedge condition, although no changes in knee kinematics were evident.

The influence of medially oriented foot orthoses has also been frequently explored in biomechanical literature. Boldt et al, examined the effects of medially wedged foot orthoses on knee and hip joint mechanics during running in females with and without patellofemoral pain syndrome [16]. The findings from this study showed that the peak knee adduction moment increased and hip adduction excursion decreased significantly when wearing medially wedged foot orthoses. Sinclair et al.,  explored the effects of medial foot orthoses on patellofemoral stress during the stance phase of running using a musculoskeletal modelling approach [17]. Their findings showed that medial foot orthoses significantly reduced peak patellofemoral stress loading at this joint during running.

Although the effects of medial/lateral foot orthoses have been explored previously, they have habitually been examined during walking in pathological patients and thus their potential prophylactic effects on the kinetics and tibiocalcaneal kinematics of running have yet to be examined. Therefore, the aim of the current investigation was to examine the effects of foot orthotic devices with a 5° medial and lateral wedge on kinetics and tibiocalcaneal kinematics the during the stance phase of running. A clinical investigation of this nature may provide further insight into the potential efficacy of wedged foot orthoses for the prevention of chronic running injuries.

Methods

Participants

Twelve male runners (age 26.23 ± 5.76 years, height 1.79 ± 0.11 cm and body mass 73.22 ± 6.87 kg) volunteered to take part in this study. All runners were free from musculoskeletal pathology at the time of data collection and were not currently taking any medications. The participants provided written informed consent in accordance with the principles outlined in the Declaration of Helsinki. The procedure utilized for this investigation was approved by the University of Central Lancashire, Science, Technology, Engineering and Mathematics, ethical committee.

Orthoses

Commercially available orthotics (Slimflex Simple, High Density, Full Length, Algeos UK) were examined in the current investigation. The orthoses were made from Ethylene-vinyl acetate and had a shore A rating of 65. The orthoses were able to be modified to either a 5˚ varus or valgus configuration which spanned the full length of the device. The order that participants ran in each orthotic condition was counterbalanced.

Procedure

Participants completed five running trials at 4.0 m/s ± 5%. The participants struck an embedded piezoelectric force platform (Kistler Instruments, Model 9281CA) sampling at 1000 Hz with their right foot. Running velocity was monitored using infrared timing gates (SmartSpeed Ltd UK). The stance phase of the running cycle was delineated as the time over which > 20 N vertical force was applied to the force platform. Kinematic information was collected using an eight-camera optoelectric motion capture system with a capture frequency of 250 Hz. Synchronized kinematic and ground reaction force data were obtained using Qualisys track manager software (Qualisys Medical AB, Goteburg, Sweden).

The calibrated anatomical systems technique (CAST) was utilized to quantify tibiocalcaneal kinematics (18). To define the anatomical frames of the right foot, and shank, retroreflective markers were positioned onto the calcaneus, first and fifth metatarsal heads, medial and lateral malleoli, medial and lateral epicondyle of the femur. A carbon fiber tracking cluster was attached to the shank segment. The foot was tracked using the calcaneus, and first and fifth metatarsal markers. Static calibration trials were obtained with the participant in the anatomical position in order for the positions of the anatomical markers to be referenced in relation to the tracking clusters/markers.

Tibial accelerations were measured using an accelerometer (Biometrics ACL 300, Units 25-26 Nine Mile Point Ind. Est. Cwmfelinfach, Gwent United Kingdom) sampling at 1000 Hz. The device was attached to the tibia 0.08 m above the medial malleolus in alignment with its longitudinal axis (19). Strong adhesive tape was placed over the device and the lower leg to prevent artifact in the acceleration signal.

Processing

The running trials were digitized using Qualisys Track Manager (Qualysis, Sweden) and then exported as C3D files. Kinematic parameters were quantified using Visual 3-D software (C-Motion, USA) after the marker data was smoothed using a low-pass Butterworth 4th order zero-lag filter at a cutoff frequency of 12 Hz. Three-dimensional kinematic parameters were calculated using an XYZ cardan sequence of rotations where X represents the sagittal plane, Y represents the coronal plane and Z represents the transverse plane rotations (Sinclair et al., 2013). Trials were normalized to 100% of the stance phase then processed and averaged. In accordance with previous studies, the foot segment coordinate system was referenced to the tibial segment for ankle kinematics, whilst tibial internal rotation (TIR) was measured as a function of the tibial coordinate system in relation to the foot coordinate axes [21]. The 3-D kinematic tibiocalcaneal measures which were extracted for statistical analysis were: (1) angle at foot strike, (2) peak angle during stance and (3) relative range of motion (ROM) from footstrike to peak angle.

The tibial acceleration signal was filtered using a 60 Hz Butterworth zero lag 4th order low pass filter to prevent any resonance effects on the acceleration signal. Peak tibial acceleration (g) was defined as the highest positive axial acceleration peak measured during the stance phase. Average tibial acceleration slope (g/s) was quantified by dividing peak tibial acceleration by the time taken from footstrike to peak tibial acceleration and instantaneous tibial acceleration slope (g/s) was quantified as the maximum increase in acceleration between frequency intervals. From the force platform all parameters were normalized by dividing the net values by body weight. Instantaneous loading rate (BW/s) was calculated as the maximum increase in vertical force between adjacent data points.

Statistical analyses

Means, standard deviations and 95 % confidence intervals were calculated for each outcome measure for all orthotic conditions. Differences in kinetic and tibiocalcaneal kinematic parameters between orthoses were examined using one-way repeated measures ANOVAs, with significance accepted at the P≤0.05 level. Effect sizes were calculated using partial eta2 (pη2). Post-hoc pairwise comparisons were conducted on all significant main effects. The data was screened for normality using a Shapiro-Wilk which confirmed that the normality assumption was met. All statistical actions were conducted using SPSS v23.0 (SPSS Inc., Chicago, USA).

Results

Tables 1-3 and Figure 1 present differences in kinetics and tibiocalcaneal kinematics as a function of the different orthoses. The results indicate that the experimental orthoses significantly affected orthoses tibiocalcaneal kinematic parameters.

Medial Lateral No-orthotic
Mean SD 95% CI (Lower) 95% CI (Upper) Mean SD 95% CI (Lower) 95% CI (Upper) Mean SD 95% CI (Lower) 95% CI (Upper)
Coronal plane (+ = inversion & – = eversion)
 Angle at footstrike (°) -3.98 5.65 -7.57 -0.39 -3.77 5.64 -7.35 -0.19 -0.66 5.91 -4.41 3.09
 Peak eversion (°) -10.75 5.7 -14.38 -7.13 -14.11 6.48 -18.22 -9.99 -12.37 5.43 -15.82 -8.92
 Relative ROM (°) 6.77 2.78 5.00 8.54 10.34 3.44 8.15 12.53 11.71 3.26 9.64 13.78
Transverse plane (+ = external & – = internal)
 Angle at footstrike (°) -11.78 2.72 -13.51 -10.05 -15.01 2.81 -16.80 -13.22 -14.41 2.97 -16.29 -12.52
 Peak rotation (°) -6.80 3.10 -8.78 -4.83 -5.6 3.94 -8.10 -3.09 -5.05 3.33 -7.17 -2.93
 Relative ROM (°) 4.97 0.86 4.43 5.52 9.41 1.33 8.56 10.26 9.35 1.44 8.44 10.27

Table 1 Ankle kinematics (mean, SD & 95% CI) in the coronal and transverse planes as a function of the different orthotic conditions.

Medial Lateral No-orthotic
Mean SD 95% CI (Lower) 95% CI (Upper) Mean SD 95% CI (Lower) 95% CI (Upper) Mean SD 95% CI (Lower) 95% CI (Upper)
Transverse plane (+ =  internal & – =external)
 Angle at footstrike (°) 8.57 3.16 6.56 10.57 9.74 4.01 7.20 12.29 6.51 3.98 3.98 9.04
 Peak TIR (°) 13.54 4.28 10.82 16.27 15.89 5.65 12.30 19.48 13.98 4.58 11.07 16.89
 Relative ROM (°) 4.98 2.68 3.28 6.68 6.14 3.54 3.89 8.39 7.47 3.75 5.09 9.85

Table 2 Tibial internal rotation parameters (mean, SD & 95% CI) as a function of the different orthotic conditions.

Medial Lateral No-orthotic
Mean SD 95% CI (Lower) 95% CI (Upper) Mean SD 95% CI (Lower) 95% CI (Upper) Mean SD 95% CI (Lower) 95% CI (Upper)
Peak tibial acceleration (g) 9.83 4.50 6.98 12.69 9.97 4.88 6.87 13.07 9.41 4.76 6.38 12.44
Average tibial acceleration slope (g/s) 362.73 196.31 238.01 487.46 367.37 219.63 227.83 506.91 369.52 257.85 205.69 533.35
Instantaneous tibial acceleration slope (g/s) 866.20 459.40 574.31 1158.09 867.71 554.16 515.61 1219.81 776.85 529.86 440.20 1113.51
Instantaneous load rate (BW/s) 156.17 58.72 118.86 193.48 161.77 71.57 116.30 207.25 134.49 44.62 106.14 162.84

Table 3 Kinetic and tibial acceleration parameters (mean, SD & 95% CI) as a function of the different orthotic conditions.

Figure 1 Tibiocalcaneal kinematics as a function of the different orthotic conditions; a= ankle coronal plane angle, b= ankle transverse plane angle & c = tibial internal rotation, (black = lateral, dash = medial & grey = no-orthotic), (IN = inversion, EXT = external & INT = internal).

Kinetics and tibial accelerations

No significant (P>0.05) differences in kinetics/tibial acceleration parameters were observed between orthotic conditions.

Tibiocalcaneal kinematics

In the coronal plane a significant main effect (F (2, 22) = 25.58, P<0.05, pη2 = 0.70) was found for the magnitude of peak eversion. Post-hoc pairwise comparisons showed that peak eversion was significantly larger in the lateral in relation to the medial (P=0.0000007) and no-orthotic (P=0.01) conditions. In addition, it was also revealed that peak eversion was significantly greater in the no-orthotic (P=0.008) in comparison to the medial orthotic condition. In addition, a significant main effect (F (2, 22) = 25.58, P<0.05, pη2 = 0.74) was noted for relative eversion ROM. Post-hoc pairwise comparisons showed that relative eversion ROM was significantly larger in the lateral (P=0.0000006) and no-orthotic (P=0.00001) in relation to the medial condition.

In the transverse plane a significant main effect (F (2, 22) = 116.11, P<0.05, pη2 = 0.91) was noted for relative transverse plane ankle ROM. Post-hoc pairwise comparisons showed that relative transverse plane ankle ROM was significantly larger in the lateral (P=0.0000001) and no-orthotic (P=0.0000008) in relation to the medial condition.

In addition, a significant main effect (F (2, 22) = 5.99, P<0.05, pη2 = 0.36) was found for the magnitude of peak TIR. Post-hoc pairwise comparisons showed that peak TIR was significantly larger in the lateral in relation to the medial (P=0.007) and no-orthotic (P=0.025) conditions. Finally, a significant main effect (F (2, 22) = 7.55, P<0.05, pη2 = 0.41) was noted for relative TIR ROM. Post-hoc pairwise comparisons showed that relative TIR ROM was significantly larger in the lateral (P=0.04) and no-orthotic (P=0.007) in relation to the medial condition.

Discussion

The aim of the current investigation was to examine the effects of foot orthotic devices with a 5° medial and lateral wedge on kinetics and tibiocalcaneal kinematics the during the stance phase of running. This is, to the authors’ knowledge, the first investigation to concurrently examine the influence of different orthotic wedge configurations on the biomechanics of running. An investigation of this nature may, therefore, provide further insight into the potential prophylactic efficacy of wedged foot orthoses for the prevention of chronic running injuries.

The current study importantly confirmed that no significant differences in impact loading or axial tibial accelerations were evident as a function of the experimental orthotic conditions. This observation opposes those of Sinclair et al., Laughton et al. and Dixon, who demonstrated that foot orthoses significantly reduced the magnitude of axial impact loading during the stance phase of running [5,7,8]. However, the findings are in agreement with those noted by Butler et al,  who similarly observed that the magnitude of axial impact loading did not differ significantly whilst wearing rigid orthoses [6]. Although not all of the aforementioned investigations have published hardness ratings, at a shore A grade of 65 it is likely that the orthoses examined in the current explanation were more rigid than those utilized by Sinclair et al., Laughton et al. and Dixon [5,7,8]. It is proposed that the divergence between investigations relates to differences in hardness characteristics of the experimental orthoses. The magnitude of impact loading is governed by the rate of change in momentum of the decelerating limb as the foot strikes the ground [22]; as such, it appears that the orthoses examined in this analysis were not sufficiently compliant to provide any reduction in impact loading.

Of further importance to the current investigation is that the medial orthoses significantly reduced eversion and TIR parameters in relation to the lateral and no-orthotic conditions. It is likely that this observation relates to the mechanical properties of the medial wedge which is designed specifically to rotate the foot segment into a more inverted position. This finding has potential clinical significance as excessive rearfoot eversion and associated TIR parameters are implicated in the etiology of a number of overuse injuries such as tibial stress syndrome, plantar fasciitis, patellofemoral syndrome and iliotibial band syndrome [23-25]. This observation therefore suggests that medial orthoses may be important for the prophylactic attenuation of chronic running related to excessive eversion/ TIR.

The findings from the current study importantly show that whilst lateral orthoses are effective in attenuating pain symptoms [9] and reducing the magnitude of the external knee adduction moment [13-15] in patients with medial compartment tibiofemoral osteoarthritis, they may concurrently place runners at risk from chronic pathologies distinct from the medial aspect of the tibiofemoral joint. It appears based on the findings from the current investigation that caution should be exercised when prescribing lateral wedge orthoses without a thorough assessment of the runners’ individual characteristics.  

A limitation, in relation to the current investigation, is that only the acute effects of the wedged insoles were examined. Therefore, although the medial orthoses appear to prophylactically attenuate tibiocalcaneal risk factors linked to the etiology of injuries, it is currently unknown whether this will prevent or delay the initiation of injury symptoms. Furthermore, the duration over which the orthoses would need to be utilized in order to mediate a clinically meaningful change in patients is also not currently known. A longitudinal examination of medial/lateral orthoses in runners would therefore be of practical and clinical relevance in the future. A further potential limitation is that only male runners were examined as part of the current investigation. Females are known to exhibit distinct tibiocalcaneal kinematics when compared to male recreational runners, with females being associated with significantly greater eversion and TIR parameters compared to males [26]. Furthermore, females are renowned for being at increased risk from tibiofemoral joint degeneration in comparison to males [27], and experimental findings have shown that degeneration may also be more prominent at different anatomical aspects of the knee in females in relation to males [28]. This suggests that the requirements of females, in terms of wedged orthotic intervention, may differ from those of male runners, thus it would be prudent for future biomechanical investigations to repeat the current study using a female sample.

In conclusion, despite the fact that the biomechanical effects of wedged foot orthoses have been examined previously, current knowledge with regards to the effects of medial and lateral orthoses on the kinetics and tibiocalcaneal kinematics of running have yet to be explored. This study adds to the current literature in the field of biomechanics by giving a comprehensive comparative examination of kinetic and tibiocalcaneal kinematic parameters during the stance phase of running whilst wearing medial and lateral orthoses. The current investigation importantly showed that medial orthoses significantly attenuated eversion and TIR parameters in relation to the lateral and no-orthotic conditions. The findings from this study indicate therefore that medial orthoses may be effective in attenuating tibiocalcaneal kinematic risk factors linked to the etiology of chronic pathologies in runners.

References

  1. Lee, D.C., Pate, R.R., Lavie, C.J., Sui, X., Church, T.S., Blair S.N. (2014). Leisure-time running reduces all-cause and cardiovascular mortality risk. Journal of the American College of Cardiology. 64, 472-481.
  2. van Gent, B.R., Siem, D.D., van Middelkoop, M., van Os, T.A., Bierma-Zeinstra, S.S., Koes, B.B. (2007). Incidence and determinants of lower extremity running injuries in long distance runners: a systematic review. British Journal of Sports Medicine. 41, 469-480.
  3. McMillan, A., Payne, C. (2008). Effect of foot orthoses on lower extremity kinetics during running: a systematic literature review. Journal of Foot and Ankle Research. 13, 1-13.
  4. Mills, K., Blanch, P., Chapman, A. R., McPoil, T. G., Vicenzino, B. (2010). Foot orthoses and gait: a systematic review and meta-analysis of literature pertaining to potential mechanisms. British Journal of Sports Medicine, 44, 1035-1046.
  5. Sinclair, J., Isherwood, J., Taylor, P.J. (2014). Effects of foot orthoses on kinetics and tibiocalcaneal kinematics in recreational runners. Foot and Ankle Online Journal, 7, 3-11.
  6. Butler, R. J., Davis, I. M., Laughton, C. M., Hughes, M. (2003). Dual-function foot orthosis: effect on shock and control of rearfoot motion. Foot & ankle international, 24, 410-414.
  7. Laughton, C. A., Davis, I. M., Hamill, J. (2003). Effect of strike pattern and orthotic intervention on tibial shock during running. Journal of Applied Biomechanics, 19, 153-168.
  8. Dixon, S.J. (2007). Influence of a commercially available orthotic device on rearfoot eversion and vertical ground reaction force when running in military footwear. Military medicine, 172, 446-450.
  9. Parkes, M. J., Maricar, N., Lunt, M., LaValley, M. P., Jones, R. K., Segal, N. A., Felson, D. T. (2013). Lateral wedge insoles as a conservative treatment for pain in patients with medial knee osteoarthritis: a meta-analysis. JAMA, 310, 722-730.
  10. Reilly, K. A., Barker, K. L., Shamley, D. (2006). A systematic review of lateral wedge orthotics-how useful are they in the management of medial compartment osteoarthritis?. The Knee, 13, 177-183.
  11. Rafiaee, M., Karimi, M. T. (2012). The effects of various kinds of lateral wedge insoles on performance of individuals with knee joint osteoarthritis. International Journal of Preventive Medicine, 3, 693-698.
  12. Birmingham, T.B., Hunt, M.A., Jones, I.C., Jenkyn, T.R., Giffin, J.R. (2007). Test–retest reliability of the peak knee adduction moment during walking in patients with medial compartment knee osteoarthritis. Arthritis Care & Research. 57, 1012-1017.
  13. 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. American Journal of Physical Medicine & Rehabilitation, 84, 858-864.
  14. Butler, R. J., Marchesi, S., Royer, T., Davis, I. S. (2007). The effect of a subject‐specific amount of lateral wedge on knee mechanics in patients with medial knee osteoarthritis. Journal of Orthopaedic Research, 25, 1121-1127.
  15. 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. American Journal of Physical Medicine & Rehabilitation, 83, 273-278.
  16. Boldt, A.R., Willson, J.D., Barrios, J.A., Kernozek, T.W. (2013). Effects of medially wedged foot orthoses on knee and hip joint running mechanics in females with and without patellofemoral pain syndrome. Journal of Applied Biomechanics. 29, 68-77.
  17. Sinclair, J., Vincent, H., Selfe, J., Atkins, S., Taylor, P.J., Richards, J. (2015). Effects of foot orthoses on patellofemoral load in recreational runners. Foot and Ankle Online Journal, 8, 5-12.
  18. Cappozzo, A., Catani, F., Leardini, A., Benedeti, M.G., Della, C.U. (1995). Position and orientation in space of bones during movement: Anatomical frame definition and determination. Clinical Biomechanics, 10, 171-178.
  19. Sinclair, J., Bottoms, L., Taylor, K., Greenhalgh, A. (2010). Tibial shock measured during the fencing lunge: the influence of footwear. Sports Biomechanics, 9, 65-71.
  20. Sinclair, J., Taylor, P.J., Edmundson, C.J., Brooks, D., Hobbs, S.J. (2013). Influence of the helical and six available Cardan sequences on 3D ankle joint kinematic parameters. Sports Biomechanics, 11, 430-437.
  21. Eslami, M., Begon, M., Farahpour, N., Allard, P. (200). Forefoot–rearfoot coupling patterns and tibial internal rotation during stance phase of barefoot versus shod running. Clinical Biomechanics, 22, 74-80.
  22. Whittle, M.W. (1999). Generation and attenuation of transient impulsive forces beneath the foot: a review. Gait & posture, 10, 264-267.
  23. Viitasalo, J.T., Kvist, M. (1983). Some biomechanical aspects of the foot and ankle in athletes with and without shin splints. The American Journal of Sports Medicine, 11, 125-130.
  24. Lee, S.Y., Hertel, J., Lee, S.C. (2010). Rearfoot eversion has indirect effects on plantar fascia tension by changing the amount of arch collapse. The Foot, 20, 64-70.
  25. Barton, C. J., Levinger, P., Menz, H. B., Webster, K. E. (2009). Kinematic gait characteristics associated with patellofemoral pain syndrome: a systematic review. Gait & posture, 30, 405-416.
  26. Sinclair, J., Taylor, P. J. (2014). Sex differences in tibiocalcaneal kinematics. Human Movement, 15, 105-109.
  27. Hame, S.L., Alexander, R.A. (2013). Knee osteoarthritis in women. Current Reviews in Musculoskeletal Medicine. 6, 182-187.
  28. Hanna, F.S., Teichtahl, A.J., Wluka, A.E., Wang, Y., Urquhart, D.M., English, D.R., Cicuttini, F.M. (2009). Women have increased rates of cartilage loss and progression of cartilage defects at the knee than men: a gender study of adults without clinical knee osteoarthritis. Menopause. 16, 666-670.

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The effects of CrossFit and minimalist footwear on Achilles tendon kinetics during running

Posted on December 31, 2016 | Comments Off on The effects of CrossFit and minimalist footwear on Achilles tendon kinetics during running

by Jonathan Sinclair1, and Benjamin Sant1pdflrg

The Foot and Ankle Online Journal 9 (4): 2

The aim of the current investigation was to comparatively assess the influence of barefoot, CrossFit, minimalist and conventional footwear on the loads experienced by the Achilles tendon during running. Twelve male runners (27.81 ± 7.02 years, height 1.77 ± 0.11 cm and body mass 76.22 ± 7.04 kg) ran at 4.0 m·s-1 in each of the four footwear conditions. Achilles tendon forces were calculated using a musculoskeletal modelling approach allowing the magnitudinal and temporal aspects of the Achilles tendon force to be quantified. Differences between footwear were examined using one-way repeated measures ANOVA. The results showed the peak Achilles tendon force was significantly larger when running barefoot (5.81 ± 1.21) and in minimalist footwear (5.64 ± 1.03 BW) compared to conventional footwear (5.15 ± 1.05 BW). In addition it was revealed that Achilles tendon impulse was significantly larger when running barefoot (0.77 ± 0.22 BW.s) and in minimalist footwear (0.72 ± 0.16 BW.s) in comparison to both conventional footwear (0.64 ± 0.15 BW.s). Given the proposed association between high Achilles tendon forces and tendon degradation, the outcomes from the current investigation indicate that CrossFit athletes who select barefoot and minimalist footwear for their running training may be at increased risk from Achilles tendon pathology in comparison to conventional footwear conditions.

Keywords: Footwear, Achilles tendon, running, CrossFit

ISSN 1941-6806
doi: 10.3827/faoj.2016.0904.0002

1 – Centre for Applied Sport and Exercise Sciences, School of Sport and Wellbeing, College of Health & Wellbeing, University of Central Lancashire, Lancashire, UK.
* – Corresponding author: jksinclair@uclan.ac.uk

CrossFit represents a relatively new activity associated with aerobic exercises, calisthenics, and Olympic weightlifting [1]. CrossFit as a discipline has expanded to become an international sport which has been linked to significant gains in aerobic and anaerobic fitness [1]. Given the novelty of CrossFit in relation to more established sports it has received a paucity of published attention in the sports science and strength and conditioning literature.

A key feature of CrossFit training is aerobic conditioning and the manner in which this is examined during competition is via distance running events. Engagement in distance running mediates numerous physiological benefits but it is known to be associated with a high rate of chronic pathologies, with around 70 % of runners experiencing an injury injured during the course of a year [2,3]. Shorten proposes that athletic footwear with suitable mechanical features may be able to manage the incidence of chronic running related injuries [4].

CrossFit athletes are able to select from a wide range of different footwear conditions with distinct design characteristics. There has been no peer reviewed research which has examined the biomechanical influence of different footwear available to CrossFit athletes. CrossFit specific footwear represents a hybrid footwear designed to incorporate the stability characteristics of a weightlifting shoe with the cushioning and flexibility of a running trainer. Currently, there is a trend for CrossFit athletes to opt to train and compete either barefoot or minimalist footwear in lieu of traditional footwear options, although the efficacy of barefoot and minimalist footwear is not yet fully established.

The effects of different footwear on the loads experienced by the Achilles tendon have been examined previously. Sinclair examined the effect running barefoot had on minimalist and conventional footwear on Achilles tendon kinetics during the stance phase of running [5]. The findings showed that peak Achilles tendon kinetics were significantly larger when running barefoot and in minimalist footwear. Similarly Sinclair et al, examined the effects of minimalist, maximalist and conventional footwear on the loads borne by the Achilles tendon during running[6] . Their findings confirmed that peak Achilles tendon force and Achilles tendon impulse were significantly larger in minimalist footwear in relation to the conventional and maximalist conditions. Currently there are no published scientific investigations regarding the effects of barefoot, CrossFit, minimalist and conventional footwear on the loads experienced by the Achilles tendon.    

Therefore the aim of the current study was to comparatively examine the influence of barefoot, CrossFit, minimalist and conventional footwear on the loads experienced by the Achilles tendon during the stance phase of running. Given that running activities are associated with a high incidence of chronic Achilles tendon pathologies, the current investigation may deliver key information to CrossFit athletes concerning the selection of suitable footwear.

Methods

Participants

Thirteen male participants took part in this investigation. All uninjured at the time of data collection and written informed consent was obtained. The mean and standard deviation (SD) characteristics of the participants were: age 27.81 ± 7.02 years, height 1.77 ± 0.11 cm and body mass 76.22 ± 7.04 kg. The research design utilized for this investigation was approved by the University of Central Lancashire, Science, Technology, Engineering and Mathematics, ethical committee. 

Procedure

Participants ran at 4.0 m·s-1 (±5%), while striking an embedded piezoelectric force platform (Kistler, Kistler Instruments Ltd., Alton, Hampshire) which sampled at 1000 Hz. Participants struck the platform with their right foot which was used for analysis. Running velocity was monitored using infrared timing gates (Newtest, Oy Koulukatu, Finland). The stance phase was delineated as the duration over which 20 N or greater of vertical force was applied to the force platform. Runners completed five trials in each footwear condition. The order that participants ran in each footwear condition was randomized. Kinematics and ground reaction forces data were synchronously collected. Kinematic data was captured at 250 Hz via an eight camera motion analysis system (Qualisys Medical AB, Goteburg, Sweden). Dynamic calibration of the motion capture system was performed before each data collection session.

Lower extremity segments were modelled in 6 degrees of freedom using the calibrated anatomical systems technique [7]. To define the segment coordinate axes of the foot and shank, retroreflective markers were placed unilaterally onto the 1st metatarsal, 5th metatarsal, calcaneus, medial and lateral malleoli, medial and lateral epicondyles of the femur. A carbon fiber tracking cluster was positioned onto the shank segment and the foot was tracked using the 1st metatarsal, 5th metatarsal and calcaneus markers. The center of the ankle joint was delineated as the midpoint between the malleoli markers[8] . Static calibration trials were obtained allowing for the anatomical markers to be referenced in relation to the tracking markers/ clusters. The Z (transverse) axis was oriented vertically from the distal segment end to the proximal segment end. The Y (coronal) axis was oriented in the segment from posterior to anterior. Finally, the X (sagittal) axis orientation was determined using the right hand rule and was oriented from medial to lateral.

Processing

Dynamic trials were digitized using Qualisys Track Manager in order to identify anatomical and tracking markers then exported as C3D files to Visual 3D (C-Motion, Germantown, MD, USA). Ground reaction force and marker trajectories were smoothed using cut-off frequencies of 50 and 12 Hz using a low-pass Butterworth 4th order zero lag filter. All data were normalized to 100% of the stance phase then processed trials were averaged. Joint kinetics were computed using Newton-Euler inverse-dynamics. To quantify net joint moments anthropometric data, ground reaction forces and angular kinematics were used.  

Achilles tendon force (BW) was determined using a musculoskeletal modelling approach. This model has been used previously to resolve differences in Achilles tendon force between different footwear [5, 6]. Achilles tendon force was quantified as a function of the plantarflexion moment (PFM) divided by the Achilles tendon moment arm (MA). The moment arm was quantified as a function of the ankle sagittal plane angle (ak) using the procedure described by Self and Paine [9]:

Achilles tendon force = PFM / MA

MA = -0.5910 + 0.08297 ak – 0.0002606 ak2

Average Achilles tendon load rate was quantified as the Achilles tendon force divided by the time over which the peak force occurred. Instantaneous Achilles tendon load rate was also determined as the peak increase in Achilles tendon force between adjacent data points. In addition to this Achilles tendon force, impulse  was quantified during running by multiplying the Achilles tendon force estimated during the stance phase by the stance time.

Experimental footwear

The footwear used during this study consisted of conventional footwear (New Balance 1260 v2), minimalist (Vibram five-fingers, ELX) and CrossFit (Reebok CrossFit CR) footwear, (shoe size 8–10 in UK men’s sizes).

Analyses

Means and standard deviations were calculated for all footwear conditions. Differences in Achilles tendon parameters between footwear were examined using one-way repeated measures ANOVAs, with significance accepted at the P≤0.05 level. Effect sizes were calculated using partial eta2 (pη2). Post-hoc pairwise comparisons were conducted on all significant main effects. The data was screened for normality using a Shapiro-Wilk which confirmed that the normality assumption was met. All statistical actions were conducted using SPSS v22.0 (SPSS Inc., Chicago, USA).

Results

Table 1 and Figure 1 present the Achilles tendon loads during the stance phase of running, as a function of the different experimental footwear. The results indicate that the experimental footwear significantly influenced Achilles tendon force parameters.

Barefoot CrossFit Conventional Minimalist
Mean SD Mean SD Mean SD Mean SD
Peak Achilles tendon force (BW) 5.81 1.21 5.50 1.32 5.15 1.05 5.64 1.03
Time to peak Achilles tendon force (s) 0.13 0.02 0.14 0.02 0.15 0.02 0.14 0.02
Achilles tendon average load rate (BW/s) 45.54 12.76 42.37 14.44 35.76 10.49 40.84 9.07
Achilles tendon instantaneous load rate (BW/s) 128.84 42.10 153.23 51.56 115.45 40.08 136.21 25.93
Achilles tendon impulse (BW.s) 0.77 0.22 0.69 0.20 0.64 0.15 0.72 0.16

Table 1 Achilles tendon forces as a function of footwear.

fig1

Figure 1 Achilles tendon forces during the stance phase as a function of footwear (black = barefoot, dash = minimalist, grey = conventional, grey dot = CrossFit).

A main effect (P<0.05, pη2 = 0.21) was shown for the magnitude of peak Achilles tendon load. Post-hoc pairwise comparisons showed that peak Achilles tendon force was significantly larger in the barefoot (P=0.01) and minimalist (P=0.04) conditions in relation to conventional footwear. A main effect (P<0.05, pη2 = 0.43) was shown for the time to peak Achilles tendon load. Post-hoc pairwise comparisons showed that time to peak Achilles tendon force was significantly larger in the barefoot (P=0.001) and minimalist (P=0.007) conditions in relation to conventional footwear. In addition time to peak Achilles tendon force was significantly shorter in the barefoot condition (P=0.007) in relation to the CrossFit footwear.  In addition a main effect (P<0.05, pη2 = 0.29) was evident for average Achilles tendon load rate. Post-hoc analysis showed that average load rate was significantly larger in the barefoot (P=0.004), CrossFit (P=0.04) and minimalist (P=0.02) conditions in relation to the conventional footwear. A main effect (P<0.05, pη2 = 0.25) was found for instantaneous Achilles tendon load rate. Post-hoc pairwise comparisons showed that instantaneous Achilles tendon load rate was significantly larger in the barefoot (P=0.01), CrossFit (P=0.003) and minimalist (P=0.01) conditions in relation to the conventional footwear. Finally, a main effect (P<0.05, pη2 = 0.34) was observed for Achilles tendon impulse. Post-hoc analyses indicated that Achilles tendon impulse was significantly larger in the barefoot (P=0.007) and minimalist (P=0.04) conditions in relation to conventional footwear.

Discussion

The aim of the current investigation was to comparatively examine the influence of barefoot, CrossFit, minimalist and conventional footwear on the loads experienced by the Achilles tendon during running. To the authors knowledge the current study represents the first comparative examination of Achilles tendon kinetics when running in these specific footwear conditions.

The key observation from the current study is that Achilles tendon force parameters were significantly larger in the barefoot and minimalist conditions in relation to the conventional running shoes. This observation is in agreement with those of Sinclair [5] and Sinclair et al. [6] who similarly noted that barefoot and minimalist footwear conditions were associated with significant increases in Achilles tendon kinetics in relation to more substantial running footwear. This observation may provide important clinical information with regards to the etiology of Achilles tendon pathologies as a function of running activities in CrossFit athletes. The initiation and progression of Achilles tendonitis is mediated by excessive tendon loads that are applied without sufficient cessation between activities [10]. Mechanisms of tendon loading that are above the systematic threshold for collagen related synthesis lead ultimately to degradation of the collagen network as the rate of resynthesis is not able to keep pace with the rate of breakdown [11]. Therefore the findings from the current investigation indicate that running barefoot and in minimalist footwear may place CrossFit athletes performing running activities at a greater risk from Achilles tendon pathology.

The specific findings from the current study may be explained by taking into account the effects of running barefoot and in minimalist footwear on the sagittal plane mechanics of the ankle joint. When running barefoot and wearing minimalist footwear, runners adopt a more plantarflexed foot position throughout the stance phase in relation to more structured running shoes [12, 13].  Increased ankle plantarflexion serves to reduce the length of moment arm of the Achilles tendon [9]. If the moment arm is shortened, this mediates an increase in the load which must be borne by the tendon when running barefoot and in minimalist footwear.

Research which has examined the influence of different footwear condition on the loads borne by the Achilles tendon during the stance phase of running, habitually examines only the peak forces experienced per footfall. Because running barefoot and in minimalist footwear mediates alterations in stance times and step frequencies, the time integral of loads experienced by the Achilles tendon are not quantified. The current study addresses this by quantifying the impulse experienced by the Achilles tendon during the stance phase which is a reflection of both the load experienced and the time interval of the load. The findings from the current investigation in relation to the Achilles tendon impulse mirror those of Sinclair et al., in that barefoot and minimalist footwear were associated with significantly larger impulse in relation to conventional running shoes [6]. This therefore further supports the notion outlines above that running barefoot and in minimalist footwear may increase the likelihood of experiencing an Achilles tendon injury compared to conventional running shoes.   

In conclusion, although differences in Achilles tendon loading as a function of different footwear conditions has been examined previously, the current knowledge with regards to the effects of minimalist, barefoot, CrossFit and conventional footwear on Achilles tendon forces is limited. As such the present study therefore adds to the current knowledge by providing a comprehensive evaluation of Achilles tendon load parameters when running in minimalist, barefoot, CrossFit and conventional footwear. On the basis Achilles tendon load and impulse parameters were shown to be significantly greater when running barefoot and in minimalist footwear, the outcomes from the current investigation indicate that CrossFit athletes who select barefoot and minimalist footwear for their running training may be at increased risk from Achilles tendon pathology in comparison to conventional footwear conditions.

References

  1. Weisenthal, B. M., Beck, C. A., Maloney, M. D., DeHaven, K. E., & Giordano, B. D. (2014). Injury rate and patterns among CrossFit athletes. Orthopaedic Journal of Sports Medicine (In press).
  2. Taunton, JE, Ryan, MB, Clement, DB, McKenzie, DC, Lloyd-Smith, DR, Zumbo, BD. A retrospective case-control analysis of 2002 running injuries. Br J Sp Med. 2002; 36: 95-101. doi: 10.1136/bjsm.36.2.95  
  3. van Gent, R, Siem DD, van Middelkoop M, van Os TA, Bierma-Zeinstra SS, Koes, BB. Incidence and determinants of lower extremity running injuries in long distance runners: a systematic review. British Journal of Sports Medicine 2007: 41: 469-480. http://www.ncbi.nlm.nih.gov/pubmed/17473005
  4. Shorten, MA. Running shoe design: protection and performance. pp 159-169 in Marathon Medicine (Ed. D. Tunstall Pedoe). 2000; London, Royal Society of Medicine.
  5. Sinclair J. Effects of barefoot and barefoot inspired footwear on knee and ankle loading during running. Clinical Biomechanics 2014; 29: 395-399. http://www.ncbi.nlm.nih.gov/pubmed/24636307
  6. Sinclair, J., Richards, J., & Shore, H. (2015). Effects of minimalist and maximalist footwear on Achilles tendon load in recreational runners. Comparative Exercise Physiology, 11(4), 239-244.
  7. Cappozzo A, Catani F, Leardini A, Benedeti MG, Della CU. Position and orientation in space of bones during movement: Anatomical frame definition and determination. Clinical Biomechanics 1995; 10: 171-178. http://www.ncbi.nlm.nih.gov/pubmed/11415549
  8. Graydon, R, Fewtrell, D, Atkins, S, Sinclair, J. The test-retest reliability of different ankle joint center location techniques. Foot Ankle Online J. 2015; 8: 1-11. doi: 10.3827/faoj.2015.0801.0011
  9. Self, BP, Paine, D. Ankle biomechanics during four landing techniques. Medicine & Science in Sports & Exercise 2001; 33: 1338–1344.
  10. Selvanetti, ACM, Puddu, G. Overuse tendon injuries: basic science and classification. Operative Techniques in Sports Medicine 1997; 5: 110–17. http://www.sciencedirect.com/science/article/pii/S1060187297800317
  11. Kirkendall, DT, Garrett W.E. Function and biomechanics of tendons. Scandinavian. Journal of Medicine & Science in Sports 1997; 7: 62–66. http://www.ncbi.nlm.nih.gov/pubmed/9211605
  12. Lieberman, D.E., Venkadesan, M., Werbel, W.A., Daoud, A.I., D’Andrea, S., Davis, I.S., & Pitsiladis, Y. (2010). Foot strike patterns and collision forces in habitually barefoot versus shod runners. Nature, 463, 531-535.
  13. Sinclair, J., Greenhalgh, A., Brooks, D., Edmundson, C. J., & Hobbs, S. J. (2013). The influence of barefoot and barefoot-inspired footwear on the kinetics and kinematics of running in comparison to conventional running shoes. Footwear Science, 5, 45-53.

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Multi-segment foot kinematics and plantar fascia strain during treadmill and overground running.

Posted on December 31, 2014 | Comments Off on Multi-segment foot kinematics and plantar fascia strain during treadmill and overground running.

By Sinclair J1, Taylor PJ2 and Vincent H1pdflrg

The Foot and Ankle Online Journal 7 (4): 4

Although physiologically beneficial, running is known to be associated with a high incidence of chronic injuries. Excessive coronal and transverse plane motions of the foot segments and strain experienced by the plantar fascia are linked to the development of a number of chronic injuries. This study examined differences in multi-segment foot kinematics and plantar fascia strain during treadmill and overground running. Twelve male recreational runners ran at 4.0 m.s-1 in both treadmill and overground conditions. Multi-segment foot kinematics and plantar fascia strain were measured using an eight-camera motion analysis system and contrasted using paired samples t-tests. The results showed that plantar fascia strain was significantly greater in the overground condition (8.23 ± 2.77) compared to the treadmill (5.53 ± 2.25). Given the proposed relationship between excessive plantar fascia strain and the etiology of injury, overground running may be associated with a higher incidence of injury although further work is necessary before causation can be confirmed.

Keywords: Running, kinematics, treadmill, plantar fascia

ISSN 1941-6806
doi: 10.3827/faoj.2014.0704.0004

Address correspondence to: Dr. Paul John Taylor, School of Psychology, University of Central Lancashire, Preston, Lancashire, PR1 2HE.
E-mail: PJTaylor@uclan.ac.uk

1 Centre for Applied Sport and Exercise Sciences, University of Central Lancashire
2 School of Psychology, University of Central Lancashire

Running using the treadmill is now a common exercise modality [1]. Recent statistics provided by runners’ world indicate that in excess of 40 million people in the US perform their running training using a treadmill. Treadmills are also useful to researchers interested in the mechanics of human gait as they allow locomotion velocity and gradient to be controlled in a controlled environment [2]. The treadmill also allows a greater number of continuous gait cycles to be captured and may thus allow more natural movement patterns to be obtained [3].

Recreational running is associated with a number of physiological benefits [4]. However, etiological analyses which have studied the prevalence of running injuries indicate that chronic injuries are extremely common, with an incidence rate of around 70 % during the course of a year [5]. A large number of retrospective and prospective analyses have investigated the mechanisms by which chronic running injuries develop [6,7,8]. Mal-alignment of the foot segment has been linked to etiology of chronic pathologies [9]. Excessive coronal and transverse plane motions of the foot segments have been associated with the progression of various pathologies such as tibial stress syndrome and anterior knee pain [10]. In addition to this, abnormal foot mechanics have also been linked to the etiology of plantar fasciitis, which affects in access of 10% or recreational runners [11].

It is currently unknown whether using the treadmill for training, compared to traditional overground running, influences runners’ susceptibility to chronic injuries. Research investigating the differences in running mechanics has habitually used a single segment foot model, thus the current understanding regarding articulations of the foot segments, linked to the potential etiology of injuries during treadmill and overground running is limited. Differences between treadmill and overground running have been examined previously in walking studies using multi-segment foot models.    Tulchin et al., [12] observed small differences in rearfoot plantarflexion during first rocker, as well as peak forefoot eversion and abduction, although all differences were <3°. These results led to the conclusion that multi-segment foot mechanics were similar between overground and treadmill walking in healthy adults.

Given the popularity of the treadmill as an exercise and research tool there has been no published information regarding the differences in multi-segment foot kinematics and plantar fascia strain during overground and treadmill running. Therefore, the aim of the current investigation was to determine whether differences exist between running on the treadmill and overground in multi-segment foot mechanics and also the strain imposed on the plantar fascia.

Methods

Participants

Twelve experienced runners took part in the current investigation. All were free from musculoskeletal pathology at the time of data collection and provided written informed consent. The mean characteristics of the participants were: Age = 24.11 ± 1.35 years, Height = 1.74 ± 0.08 m, Mass = 69.16 ± 5.67 kg. The procedure utilized for this investigation was approved by the University of Central Lancashire, ethical committee.

Procedure

Kinematic information during overground and treadmill locomotion was captured at 250 Hz via an eight-camera motion analysis system (QualisysTM Medical AB, Goteburg, Sweden). Two identical motion capture systems were used. Calibration of each system was performed before each data collection session. Calibrations producing residuals <0.85 mm and points above 4000 in all cameras were considered acceptable.

In order to model the foot segments in six degrees of freedom the calibrated anatomical systems technique was utilized for modelling and tracking segments was [13]. Circular retroreflective markers (19 mm) were placed onto specific anatomical landmarks in accordance with the foot model developed by Leardini et al., [14]. This allowed the anatomical and technical frames of the rearfoot (Rear), midfoot (Mid) and forefoot (Fore) to be delineated. To define the tibial (Tib) segment additional markers were positioned onto the medial and lateral femoral epicondyles. A rigid carbon-fibre tracking cluster consisting of four non-linear markers was also positioned onto this segment. All participants were provided with the same experimental footwear (Asics 2160; Asics UK).

In the overground condition, five trials were undertaken over a 22 m walkway (Altrosports 6mm, Altro Ltd) at a velocity of 4.0m.s-1 ±5%. The velocity of running was quantified using infra-red timing gates (SmartSpeed Ltd UK). To collect treadmill data a WoodwayTM (ELG, Germany) treadmill was utilized. Participants were allowed a habitation period of 5-min, during which they ran at the required experimental velocity prior to data collection. Five trials were also collected for the treadmill locomotion. As force information was not available from the treadmill, footstrike and toe-off were determined in both conditions using kinematic information. Based on the recommendations of Fellin et al., [15], footstrike was determined as the point at which the vertical velocity of the calcaneus marker changed from negative to positive and toe-off was delineated using the second instance of peak knee extension.

Data processing

Data were digitized using Qualisys track manager and exported to Visual 3D (C-motion, Germantown USA). Marker trajectories were smoothed at 15 Hz using a low pass non-phase shift Butterworth filter. This frequency was selected based on residual analysis [16]. Cardan angles were used to calculate 3-D articulations of the foot segments. Foot angles were calculated using and XYZ cardan sequence of rotations between the calcaneus-tibia (Cal-Tib), midfoot-calcaneus (Mid-Cal), forefoot-midfoot (For-Mid) and forefoot-calcaneus (For-Rear). Discrete 3-D kinematic parameters that were extracted for statistical analysis were 1) angle at footstrike, 2) angle at toe-off, 3) range of motion from footstrike to toe-off during stance, 4) peak angle during stance and 5) relative range of motion (representing the angular displacement from footstrike to peak angle). Plantar fascia strain was quantified in accordance with the Ferber et al., [17] recommendations by determining the distance between the 1st metatarsal and calcaneus markers and calculated as the relative position of the markers was altered. Plantar fascia strain was calculated as the peak change in length during the stance phase divided by the original length.

F1

Figure 1 Multi-segment foot kinematics during running in the a. sagittal, b. coronal and c. transverse planes as a function of different conditions (Black = overground and grey = treadmill) (DF =dorsiflexion, IN = inversion, INT = internal) (Rear = rearfoot, Mid = midfoot, Fore = forefoot, Tib = tibia).

tab1

Table 1 Rearfoot-tibial kinematics during treadmill and overground running conditions.

tab2

Table 2 Midfoot-rearfoot kinematics during treadmill and overground running conditions.

tab3

Table 3 Forefoot-midfoot kinematics during treadmill and overground running conditions.

tab4

Table 4 Forefoot-rearfoot kinematics during treadmill and overground running conditions.

Statistical analysis

Descriptive statistics (means and standard deviations) were calculated for each running condition. Differences in the outcome multi-segment foot kinematic parameters and plantar fascia strain were contrasted using paired samples t-tests with significance accepted at the p≤0.05 level [18]. Effect sizes for all significant observations were calculated using a Cohen’s D statistic. The data were screened for normality using a Shapiro-Wilk test. All statistical procedures were conducted using SPSS v22 (SPSS Inc, Chicago, USA).

Results

The results indicate that whilst the multi-segment foot kinematic waveforms measured during overground and treadmill running were quantitatively similar, significant differences were found to between the two running modalities. Figure 1 presents the 3-D multi-segment foot kinematics from the stance phase. Tables 1-5 present the results of the statistical analysis conducted on the measures of multi-segment foot kinematics.

Plantar fascia strain and stance time

Running overground was associated with significantly (t (11) = 2.71, p<0.05, D = 1.56) greater plantar fascia strain (8.23 ± 2.77) compared to running on the treadmill (5.53 ± 2.25). Stance time was shown to be significantly (t (11) = 3.45, p<0.05, D = 1.99) shorter in the overground condition (0.23 ± 0.05) compared to the treadmill (0.29 ± 0.03).

Rearfoot-tibia

Running overground was associated with significantly (t (11) = 2.37, p<0.05, D = 1.37) greater dorsiflexion at footstrike compared to running on the treadmill. In addition overground running was shown to be associated with a significantly (t (11) = 3.28, p<0.05, D = 1.89) larger sagittal plane ROM compared to the treadmill.

Midfoot-rearfoot

No significant (p>0.05) differences were observed.

Forefoot-midfoot

No significant (p>0.05) differences were observed.

Forefoot-rearfoot

No significant (p>0.05) differences were observed.

Discussion

Therefore, the aim of the current investigation was to determine whether differences exist between running on the treadmill and overground in multi-segment foot mechanics and also the strain imposed on the plantar fascia. This represents the first biomechanical examination to contrast both multi-segment foot kinematics and plantar fascia strain during treadmill and overground running.

The first key clinical observation from the current investigation is that plantar fascia strain was shown to be significantly greater in during treadmill running compared to overground. This finding may be clinically relevant with regards to the etiology and progression of plantar fasciitis, which is considered to be related to the magnitude of the strain imposed on the plantar fascia itself [19]. Currently, there is very little information regarding the different susceptibility of runners to chronic injuries during treadmill and overground running conditions. The results from the current study, therefore, provide insight into the biomechanical mechanisms that may affect injury susceptibility and suggest that running overground may place runners at increased risk from plantar fasciitis.

In addition to alterations in plantar fascia strain between conditions, a significantly different sagittal plane rearfoot angle was shown at footstrike between the two running conditions. Specifically, runners were shown to exhibit dorsiflexion during overground running and plantarflexion in the treadmill condition. This result is in agreement with the observations of Wank et al., [20] and Nigg et al., [21], who showed increased ankle plantarflexion at footstrike during treadmill running. Given the significant reduction in stance time, the change in rearfoot position relative to the tibial segment may relate to a shortened stride length. Both Chia et al., [22] and Schache et al., [2] noted reductions in stride distance during treadmill running in conjunction with increased stance times during treadmill running. This finding may also relate to a switch from a rearfoot to midfoot strike pattern, although without the presence of an instrumented treadmill with an integrated force platform, it is not possible to examine the vertical ground reaction force curves to fully ascertain this.

On the basis that increases in plantar fascia strain were noted during overground running, the results from the current may provide evidence to support the utilization of treadmill running to reduce runners’ susceptibility to injury. However it is important that these observations be contextualised by taking account the aforementioned increases in stride frequency that have been observed previously for treadmill running [2, 22]. Therefore, whilst increases in plantar fascia strain were noted for each foot contact when running overground, the amount of cumulative strain may not be altered between the two running modalities, as the total number of footfalls required to achieve required velocity is greater when running on the treadmill. There is currently no epidemiological data concerning the influence of cumulative and singular loads experienced by the musculoskeletal structures with regards to the etiology of chronic injuries. It is, therefore, strongly recommended that analyses prospectively investigate the effects of treadmill an overground running on the predisposition of recreational runners to chronic injuries.

A potential drawback to the current investigation was that the treadmill data did not feature an integrated force platform. Therefore, in addition to being unable to identify footstrike modifications this meant that footstrike and toe-off events were defined using kinematic methods. The identification of gait events using kinematic techniques has been shown to be repeatable, but they are not as accurate as the gold-standard method using force platform information [23]. Plantar fascia strain was calculated using markers placed onto the foot segment and the location of plantar fascial tissue was considered in this study to span from the calcaneus to the first metatarsal. This procedure has been adopted previously in order to model the strain experienced by the plantar fascia [17, 24] and the means strain values presented in this work closely correspond with previous values. Nonetheless, this represents a simplified technique and there is likely to be some error associated with this method. Future analyses should consider more accurate techniques, such as fluoroscopic imaging in conjunction with 3-D motion analysis, to provide a more accurate measurement of plantar fascia strain during different running conditions.

In conclusion, the current investigation adds to the current knowledge in the discipline of clinical biomechanics by providing a comprehensive evaluation of the 3-D multi-segment foot kinematics and plantar fascia strain observed when running on the treadmill and overground. This study demonstrated that plantar fascia strain was significantly reduced during treadmill running. This indicates that running on the treadmill may be associated with a reduced incidence of plantar fasciitis, although additional epidemiological research is required before specific conclusions regarding injury prevention can be made

Acknowledgements

Our thanks go to Glen Crook for his technical assistance.

References

  1. Sinclair J, Richards J, Taylor PJ, Edmundson CJ, Brooks D, Hobbs SJ. Three-dimensional kinematic comparison of treadmill and overground running. Sports Biomechanics 2013 12: 272-282. (Link)
  2. Schache AG, Blanch PD, Rath DA, Wrigley TV, Starr R, Bennell KL. A comparison of overground and treadmill running for measuring the three-dimensional kinematics of the lumbo-pelvic-hip complex. Clinical Biomechanics 2001 16: 667–680. (PubMed)
  3. Fellin RE, Manal K, Davis IS. Comparison of lower extremity kinematic curves during overground and treadmill running. Journal of Applied Biomechanics 2010 26: 407–414. (PubMed)
  4. Wen DY. Risk factors for overuse injuries in runners. Current Sports Medicine Reports 2007 6: 307–313. http://www.ncbi.nlm.nih.gov/pubmed/17883966
  5. van Mechelen W. Running injuries. Sports Medicine 1992 14: 320-335. (PubMed)
  6. Taunton JE, Ryan MB, Clement DB, McKenzie DC, Lloyd-Smith DR, Zumbo, BD. A retrospective case-control analysis of 2002 running injuries. British Journal of Sports Medicine 2002 36: 95-101. (Link)
  7. van Gent BR, Siem DD, van Middelkoop M, van Os TA, Bierma-Zeinstra SS, Koes BB. Incidence and determinants of lower extremity running injuries in long distance runners: a systematic review. British journal of sports medicine 2007 41: 469-480. (Link)
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  9. Stergiou N, Bates BT, James SL. Asynchrony between subtalar and knee joint function during running. Medicine & Science in Sports & Exercise 1999 31: 1645–55. (PubMed)
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  11. Lareau CR, Sawyer GA, Wang JH, DiGiovanni CW. Plantar and Medial Heel Pain: Diagnosis and Management. The Journal of the American Academy of Orthopaedic Surgeons 2014 22: 372–380. (PubMed)
  12. Tulchin K, Orendurff M, Karol L. A comparison of multi-segment foot kinematics during level overground and treadmill walking. Gait & posture 2010 31: 104-108. (PubMed)
  13. Cappozzo A, Catani F, Della Croce U, Leardini A. Position and orientation in space of bones during movement: anatomical frame definition and determination. Clinical Biomechanics 1995 10: 171–178. (PubMed)
  14. Leardini A, Benedetti M, Berti L, Bettinelli D, Nativo R, Giannini S. Rear-foot, mid-foot and fore-foot motion during the stance phase of gait. Gait & Posture 2007 25: 453-462. (Link)
  15. Fellin RE, Rose WC, Royer TD, Davis IS. Comparison of methods for kinematic identification of footstrike and toe-off during overground and treadmill running. Journal of Science and Medicine in Sport 2010 13: 646–650. (PubMed)
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  17. Ferber R, Benson B. Changes in multi-segment foot biomechanics with a heat-mouldable semi-custom foot orthotic device. Journal of Foot and Ankle Research 2011 4: 1-8. (PubMed)
  18. Sinclair J, Taylor PJ Hobbs SJ. Alpha level adjustments for multiple dependent variable analyses and their applicability–A review. International Journal of Sport Science and Engineering 2013 7: 17-20.
  19. Pohl MB, Messenger N, Buckley JG. Forefoot, rearfoot and shank coupling: effect of variations in speed and mode of gait. Gait & Posture 2007 25: 295-302. (PubMed)
  20. Wank V, Frick U, Schmidtbleicher D. Kinematics and electromyography of lower limb muscles in overground and treadmill running. International Journal of Sports Medicine 1998 19: 455–461. (PubMed)
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  24. Tome J, Nawoczenski DA, Flemister A, Houck J. Comparison of foot kinematics between participants with posterior tibialis tendon dysfunction and healthy controls. Journal of Orthopaedic & Sports Physical Therapy 2006 36: 635–644. (PubMed)

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Gender differences in multi-segment foot kinematics and plantar fascia strain during running

Posted on December 31, 2014 | Comments Off on Gender differences in multi-segment foot kinematics and plantar fascia strain during running

By Sinclair J1, Chockalingam N2 and Vincent H1pdflrg

The Foot and Ankle Online Journal 7 (4): 2

This study aimed to determine whether there are gender differences in multi-segment foot kinematics and plantar fascia strain during running. Fifteen male and fifteen female participants ran at 4.0- m.s-1. Multi-segment foot kinematics and plantar fascia strain were quantified using a motion capture system and compared between genders using independent samples t-tests. The results showed that plantar fascia strain was significantly greater in males (0.09 ± 0.04) compared to females (0.06 ± 0.03). Furthermore male runners (-9.72 ± 3.09) were also associated with a significantly larger peak calcaneal eversion angle compared to females (-6.03 ± 2.33). Given the proposed relationship between high levels of plantar fascia strain as well as excessive coronal plane rotations of the foot segments and the etiology of injury, it is likely that the potential risk of the developing running injuries in relation to these mechanisms is higher in males.

Keywords: Running, gender, biomechanics

ISSN 1941-6806
doi: 10.3827/faoj.2014.0704.0002

Address correspondence to: Jonathan Sinclair, Jksinclair@uclan.ac.uk

1 Division of Sport Exercise and Nutritional Sciences, School of Sport Tourism and Outdoors, University of Central Lancashire,
2 Faculty of Health Sciences, Staffordshire University.

Recreational distance running is currently an extremely popular pastime for both males and female alike [1]. Although regular running activities offer a plethora of physiological benefits [2], the susceptibility of runners to degenerative chronic injuries is also well documented [3]. In their retrospective analysis of chronic running injuries, Taunton et al [4] demonstrated that patellofemoral pain, iliotibial band syndrome, and plantar fasciitis were the most commonly experienced chronic pathologies. Female runners have been shown to be at greater risk from chronic injuries due to running in comparison to age matched males [5].

It has been frequently hypothesized, in addition to anatomical variances, that differences in lower extremity running biomechanics may be a causative mechanism that explains why females sustain different injury patterns in comparison to males [1,6,7]. Analyses investigating the prevalence of pathologies indicate females are twice as likely to sustain a chronic injury related to running compared to males [5].

Gender differences in lower extremity kinematics have been examined previously in biomechanical literature. Sinclair et al [7] determined that female runners exhibited significantly greater peak knee abduction and rotation angles in comparison to males. Similarly, Ferber et al. [6] showed a significantly greater peak hip internal rotation and adduction angle and a significantly larger peak knee abduction angle in female runners. Sinclair & Taylor [1] compared gender differences in tibiocalcaneal kinematics during the stance phase of running. They showed that peak eversion and tibial internal rotation angles were significantly greater in female runners. These studies display a clear pattern in terms of the gender differences in running biomechanics showing that differences primarily occur in the coronal and transverse planes, which may explain the increased susceptibility of female runners to chronic injuries. Each of the aforementioned investigations utilized a single segment foot model however, and did not quantify plantar fascia strain as part of their experimental protocol. Therefore, there is currently a paucity of information regarding the potential gender differences in multi-segment foot kinematics and strain experienced by the plantar fascia during running.

This study aims to determine whether there are gender differences in multi-segment foot kinematics and plantar fascia strain during the stance phase of running. A study of this nature may be beneficial to the biomechanics and clinical communities as it may provide further insight into the mechanisms by which male and female runners suffer from distinct chronic injury patterns.

Methods

Participants

Fifteen male (age 26.98 years SD 2.87, height 1.74 m SD 0.15, mass 71.66 kg SD 4.74) and fifteen female (age 24.22 years SD 2.56, height 1.68 m SD 0.16, mass 64.22 kg SD 3.79) participants volunteered to take part in this study. All were free from musculoskeletal pathology at the time of data collection and provided informed consent in written form. Ethical approval was obtained from a University ethical committee in accordance with the declaration of Helsinki.

Procedure

Participants completed five trials running at 4.0 m.s-1 ± 5%. Multi-segment foot kinematics and plantar fascia strain were quantified using an eight-camera motion analysis system (Qualisys Medical, Sweden) with a sample rate of 250 Hz. Participants struck an embedded force platform (Kistler 9281CA, Kistler Instruments, UK) sampling at 1000 Hz with their dominant foot [8]. The stance phase of running was determined as the time over which >20 N of force in the axial direction was applied to the force platform [9]. The calibrated anatomical systems technique (CAST) procedure for modelling and tracking segments was adhered to [10]. Markers were placed on anatomical landmarks in accordance with the Leardini et al. [11] foot model protocol allowing the anatomical frames of the calcaneus (Cal), midfoot (Mid), and forefoot (Fore) to be defined. Markers were positioned on the medial and lateral femoral epicondyles to allow the anatomical frame of the tibia (Tib) to be delineated and a rigid tracking cluster was also positioned on the tibia. Participants wore the same footwear throughout (Saucony Pro Grid Guide II, Saucony, USA) in sizes 5-10 men’s UK.

Data processing

Retroreflective marker trajectories were identified using Qualisys track manager and then exported to Visual 3D (C-motion, Germantown USA). Marker trajectories were filtered at 12 Hz using a low pass zero-lag Butterworth filter. Cardan angles were used to calculate 3-D articulations of the foot segments. Stance phase angles were computed using an XYZ cardan sequence of rotations between the calcaneus-tibia (Cal-Tib), midfoot-calcaneus (Mid-Cal), forefoot-midfoot (Fore-Mid), and forefoot-calcaneus (Fore-Cal). 3-D kinematic parameters which were extracted for statistical analysis were 1) angle at footstrike, 2) angles at toe-off, 3) range of motion from footstrike to toe-off during stance, 4) peak angle during stance, and 5) relative range of motion (representing the angular displacement from footstrike to peak angle). Plantar fascia strain was determined by calculating the distance between the first metatarsal and calcaneus markers and quantified as the relative position of the markers was altered. Plantar fascia strain was calculated as the change in length during the stance phase divided by the original length [12].

Statistical analysis

Descriptive statistics were calculated for both the orthotic and no-orthotic conditions. Differences in kinematic and plantar fascia strain parameters were examined using independent samples t-tests with significance accepted at the p<0.05 level. A Shapiro-Wilk test was used to screen the data for normality, it was confirmed that the normality assumption was not violated. Effect sizes for all statistical main effects were calculated using a Cohen’s D. Statistical procedures were undertaken using SPSS v21 (IBS, SPSS INC USA).

Figure1

Figure 1: Multi-segment foot kinematics during running in the a. sagittal, b. coronal and c. transverse planes as a function of gender markers (Solid=male and Dot=female) (DF=dorsiflexion, IN=inversion, INT=internal, EXT=external) (Cal=calcaneus, Mid=midfoot, Fore=forefoot, Tib=tibia).

Table1

Table 1: Cal-Tib kinematics as a function of gender. (* =significant difference)

Table2

Table 2: Mid-Cal kinematics as a function of gender. (* =significant difference)

Table3

Table 3: Fore-Mid kinematics as a function of gender. (* =significant difference)

Table4

Table 4: Fore-Cal kinematics as a function of gender. (* =significant difference)

Results

Although qualitative examination of the kinematic curves from males and females indicate that they predominately followed a similar pattern, significant differences were observed between genders. Figure 1 and Tables 1-4 present the mean multi-segment foot parameters and stance phase joint angle curves obtained as a function of gender.

Plantar fascia strain

Males (0.09 ± 0.04) were associated with a significantly (t(28)=2.55, p<0.05, D=0.96) greater plantar fascia strain compared to females (0.06 ± 0.03).

Foot kinematics

Cal-Tib

In the sagittal plane, males were shown to exhibit significantly greater dorsiflexion at footstrike (t(28)=3.35, p<0.05, D=1.27) and were also associated with a significantly larger peak dorsiflexion (t(28)=2.56, p<0.05, D=0.97) compared to females. In the coronal plane, males were shown to exhibit significantly greater eversion at footstrike (t(28)=2.35, p<0.05, D=0.89) and were also associated with a significantly larger peak eversion (t(28)=2.51, p<0.05, D=0.95) compared to females.

Mid-Cal

In the sagittal plane, females were shown to exhibit significantly greater peak dorsiflexion (t(28)=2.34, p<0.05, D=1.27) compared to males.

Fore-Mid

In the sagittal plane, females were shown to exhibit significantly greater dorsiflexion at toe-off (t(28)=2.26, p<0.05, D=0.85) and were also associated with a significantly larger peak dorsiflexion (t(28)=2.64, p<0.05, D=1.00) compared to males. In addition, females were also associated with a significantly greater range of motion (t(28)=2.88, p<0.05, D=1.09) and relative range of motion (t(28)=3.02, p<0.05, D=1.14) compared to males.

Fore-Cal

In the sagittal plane, females were shown to exhibit significantly greater dorsiflexion at toe-off (t(28)=2.34, p<0.05, D=0.88) and were also associated with a significantly larger peak dorsiflexion (t(28)=3.20, p<0.05, D=1.21) compared to males. In addition, females were also associated with a significantly greater range of motion (t(28)=3.00, p<0.05, D=1.13) and relative range of motion (t(28)=4.16, p<0.05, D=1.57) compared to males.

Discussion

The aim of the current investigation was to determine whether differences in multi-segment foot kinematics and plantar fascia strain are present between males and females. This represents the first comparative investigation to simultaneously examine multi-segment foot kinematics and plantar fascia strain in male and female runners.

The first key observation from the current investigation is that plantar fascia strain was shown to be significantly greater in male runners compared to female runners. This finding is likely to have clinical significance regarding the etiology of plantar fasciitis which is considered to be related to the magnitude of the strain imposed on the plantar fascia itself [13]. This provides further evidence to support the observations of Taunton et al. [4] who showed that males suffered a significantly higher rate of chronic injuries to the plantar fascia. The results from the current study therefore provide further insight into the biomechanical mechanisms behind the increased susceptibility of male runners to plantar fasciitis.

A further key finding from the present study is that significant gender differences were observed in the sagittal plane for all four foot articulations. Examination of the Cal-Tib articulation indicates that males were associated with a significantly greater peak dorsiflexion angle whereas at the more distal Mid-Cal, Fore-Mid, and Fore-Cal regions, larger peak dorsiflexion angles were observed in female runners. This finding opposes the results of Sinclair et al. [7] who showed using a single segment foot model that no sagittal plane differences in foot kinematics were present between genders. This observation may relate to differences in stride length characteristics between genders as males have been shown to be associated with significantly longer stride lengths than females [14]. Furthermore this finding may also be associated with differences in foot shape or structure. Wunderlich & Cavanagh [15] showed that allometrically scaled foot dimensions in runners differed between genders which could mediate alterations in foot mechanics during the stance phase.

In addition to differences in the sagittal plane, there were also significant alterations between genders in the coronal plane. Specifically, males were associated with increased peak Cal-Tib eversion. This finding disagrees with the observations of Sinclair et al. [7] who found using a single segment foot model that females were associated with significantly greater foot eversion compared to males. Given the proposed relationship between excessive coronal and transverse plane foot motions and the incidence of chronic running injuries, this finding may also have clinical relevance and suggests that males may be more susceptible to foot pathologies [13]. This observation in conjunction with the increase in plantar fascia strain opposes the current consensus in biomechanical literature, which suggests that female runners are more susceptible to chronic injury. The findings from the current study indicate that injury susceptibility may be site specific with females being more likely to suffer from chronic injuries at the hip and knee and males perhaps more susceptible to foot pathology.

There are some limitations associated with the current study. Firstly, plantar fascia strain was obtained using markers positioned onto the foot segment and the plantar fascia length itself was taken as the distance between calcaneus and first metatarsal locations. Whilst this procedure has been adopted in previous analyses to quantify the strain experienced by the plantar fascia [12], it is nonetheless a simplified practice for which there is likely to be some degree of error. Future analyses may wish to consider more direct fluoroscopic measurements in conjunction with 3-D motion capture to achieve accurate plantar fascia strain measurements. In addition, retroreflective markers placed onto the shoe in order to quantify foot articulations may also serve as a limitation as the foot is known to move relative to the shoe itself and thus the accuracy of this technique is questionable. Previous analyses have investigated the variations in foot kinematics using markers placed onto the shoe and those placed onto the skin through holes cut into the shoe itself [16]. It was demonstrated that markers positioned onto the shoe may lead to errors particularly in the coronal and transverse planes. However, because cutting holes in the footwear reduced the structural integrity of the shoe upper and also influenced the runners’ perception of the footwear, it was determined that the present technique is acceptable.

In conclusion, the current investigation provides information not previously available describing multi-segment foot kinematics and plantar fascia strain in male and female runners. Importantly, increased plantar fascia strain and peak non-sagittal angles of the Cal-Tib articulation were observed in male runners. Given the proposed relationship between high levels of plantar fascia strain as well as excessive coronal plane rotations of the foot segments and the etiology of injury, it is likely that the potential risk of the developing running injuries in relation to these mechanisms is higher in males.

Acknowledgements

Our thanks go to Robert Graydon for his technical assistance.

References

  1. Sinclair JS, Taylor PJ. Sex differences in tibiocalcaneal kinematics. Human Movement 2014 Aug;15(2):105–109. (Link)
  2. Mora S, Lee IM, Buring JE, Ridker PM. Association of physical activity and body mass index with novel and traditional cardiovascular biomarkers in women. JAMA 2006 Mar;295(12):1412-1419. (PubMed)
  3. Van Gent BR, Siem DD, van Middelkoop M, van Os TA, Bierma-Zeinstra SS, Koes B. B. Incidence and determinants of lower extremity running injuries in long distance runners: a systematic review. British Journal of Sports Medicine 2007 Aug;41(8):469-480. (PubMed)
  4. Taunton JE, Ryan MB, Clement DB, McKenzie DC, Lloyd-Smith DR, Zumbo BD. A prospective study of running injuries: the Vancouver Sun Run “In Training” clinics. British Journal of Sports Medicine 2003 Jun;37(3):239–244. (PubMed)
  5. Robinson RL, Nee RJ. Analysis of hip strength in females seeking physical therapy treatment for unilateral patellofemoral pain syndrome. JOSPT 2007 May;37(5):232–238. (PubMed)
  6. Ferber R, Davis IM, Williams DS. Gender differences in lower extremity mechanics during running. Clinical Biomechanics 2003 May;18(4):350–357. (PubMed)
  7. Sinclair J, Greenhalgh A, Edmundson CJ, Brooks D, Hobbs SJ. Gender Differences in the Kinetics and Kinematics of Distance Running: Implications for Footwear Design. International Journal of Sports Science & Engineering 2012 Jun;6(2):118–128. (Link)
  8. Sinclair J, Hobbs SJ, Taylor PJ, Currigan G, Greenhalgh A. The Influence of Different Force and Pressure Measuring Transducers on Lower Extremity Kinematics Measured During Running. Journal of Applied Biomechanics 2014 Feb;30(1):166–172. (PubMed)
  9. Sinclair J, Edmundson CJ, Brooks D, Hobbs SJ. Evaluation of kinematic methods of identifying gait events during running. International Journal of Sports Science & Engineering 2011 Sep;5(3):188–192. (Link)
  10. Cappozzo A, Catani F, Della Croce U, Leardini A. Position and orientation in space of bones during movement: anatomical frame definition and determination. Clin Biomech 1995 Jun;10(4):171–178. (PubMed)
  11. Leardini A, Benedetti M, Berti L, Bettinelli D, Nativo R, Giannini S. Rear-foot, mid-foot and fore-foot motion during the stance phase of gait. Gait Posture 2007 Mar;25(3): 453-462. (PubMed)
  12. Ferber R, Benson B. Changes in multi-segment foot biomechanics with a heat-mouldable semi-custom foot orthotic device. J Foot Ankle Res 2011;4(18):1-8. (PubMed)
  13. Pohl MB, Hamill J, Davis IS. Biomechanical and anatomic factors associated with a history of plantar fasciitis in female runners. Clinical Journal of Sport Medicine 2009 Sep;19(5):372-376. (PubMed)
  14. Elliott BC, Blanksby BA. Optimal stride length considerations for male and female recreational runners. Br J Sports Med 1979 Apr;13(1):15–18. (PubMed)
  15. Wunderlich E, Cavanagh PE. Gender differences in adult foot shape: implications for shoe design. Medicine & Science in Sport & Exercise 2001 Apr;33(4):605–611. (PubMed)
  16. Sinclair J, Greenhalgh A, Taylor PJ, Edmundson CJ, Brooks D, Hobbs SJ. Differences in tibiocalcaneal kinematics measured with skin and shoe-mounted markers. Human Movement 2013 Mar;14(1):64– 69. (Link)

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A comparison of center of pressure variables recorded during running in barefoot, minimalist footwear, and traditional running shoes in the female population

Posted on September 30, 2014 | Comments Off on A comparison of center of pressure variables recorded during running in barefoot, minimalist footwear, and traditional running shoes in the female population

by Andrew Greenhalgh1PhD, Jenny Hampson2Bsc, Peter Thain2 PhD, and Jonathan Sinclair3PhDpdflrg

The Foot and Ankle Online Journal 7 (3): 6

In recent years, barefoot running and running in minimalist footwear as opposed to running in traditional running shoes has increased in popularity. The influence of such footwear choices on center of pressure (COP) displacements and velocity variables linked to injuries is yet to be understood. The aim of this study was to investigate differences between COP variables, linked to injuries measured in barefoot running, a minimalist running shoe, and with traditional running shoes and conditions during running in a healthy female population. Seventeen healthy female participants were examined. Participants performed five footfalls in each footwear condition while running at 12km/h±10% over a pressure plate while COP variables were recorded at 500Hz. The results suggest that minimalist running shoe COP characteristics were similar to those of barefoot runners, with various significant differences reported in both groups compared to runners with the traditional running shoe.

Keywords: footwear, barefoot, running, COP, center of pressure, plantar pressure

ISSN 1941-6806
doi: 10.3827/faoj.2014.0703.0006

Address correspondence to: Andrew Greenhalgh PhD
Email: a.greenhalgh@mdx.ac.uk

1 London Sport Institute, Middlesex University, Hendon, UK
2 School of Life & Medical Sciences, University of Hertfordshire, Hatfield, UK
3 Division of Sport Exercise and Nutritional Sciences, University of Central Lancashire, Preston, UK

Following the introduction of running specific footwear, in recent years barefoot (BF) running as opposed to running in traditional running shoes (TRS) with elevated cushioned heels has increased in popularity among participants and coaches [1]. When running barefoot on roads or pathways the plantar region of the foot may be exposed to cuts and general discomfort from debris and uneven surfaces, therefore running in minimalist footwear that may allow for the change in running kinetics and kinematics observed in barefoot running compared to shod while protecting the plantar region of the feet from injury and discomfort appears to be desirable.

This has led to a rise in the popularity of barefoot inspired footwear amongst running populations and subsequent research [2]. Running barefoot does not appear to restrict athletes from competing at an elite level, with competitors winning Olympic medals in such conditions. In terms of energy cost to the runner, running barefoot appears to reducing angular inertia of the lower extremities. Research suggests minimalist shoes may also decrease oxygen consumption during running [3,4]. However, recent research suggests there is no reduction of metabolic cost when running barefoot compared to lightweight running shoes [5].

Some research suggests that wearing traditional running shoes may restrict freedom of movement and flexibility that can be achieved in comparison to barefoot running [6]. Furthermore, running barefoot compared to shod has been identified as causing adaptation in running mechanics, resulting in a more midfoot footfall in contrast to a favored heel striking movement strategy while running in traditional running shoes [2,7]. Research also suggests that such adaptations occur instantaneously with only minor changes in the lower extremity kinematics  observed in the reported knee angle after two weeks of training in minimalist footwear [8].  Such adaptations observed in barefoot running have been proposed as a mechanism by which the potentially detrimental loading imposed upon the musculoskeletal system during running may be attenuated [9–11]. Conflicting research has however reported such increases in loading of the musculoskeletal system in barefoot running compared to shod, in participants who habitually wore shoes [12,13]. Furthermore, foot injuries including stress fractures most prominently in the metatarsals have been reported in minimalist shoe runners [14]. Currently there appears to be a lack of evidence confirming the influence of barefoot running on movement strategy and injury rates [15,16].

Research identifying the influence of footwear conditions should initially focus on areas of greatest injury risk within the musculoskeletal system which research suggests is ankle ligament damage [17]. The ankle joint is unique in that the vast majority of injuries sustained across different populations are of one type;  ligament sprains [17–21]. It is worth noting that such injury rates in females [22] are higher than those of males [23].

The reason for the higher occurrence of ankle sprains while running can only be hypothesized.  Research has suggested that during running the ankle is often placed in a compromised supinated position when the athlete’s center of gravity (COG) is positioned over the lateral border of the weight bearing limb [24,25]. It has been identified that the functionally unstable ankle may be the result of proprioceptive neuromuscular deficits arising from structural damage following an injury [26–29].

Various kinetic and kinematic variables have been investigated to compare differences between barefoot and shod conditions.  However there is a paucity of research investigating the differences in center of pressure (COP) variables between the conditions [16]. Plantar COP velocities and displacements measured during running have been identified as indicators of exercise induced lower leg injuries [30,31]. As such, identifying characteristics of the COP have been identified as suitable reference points for studying the dynamics of the rearfoot and foot function [31,32] and to identify differences in footwear conditions [33]. Studies analyzing the gait of those individuals with functional unstable ankles have identified a tendency for a laterally situated COP on initial foot contact with a greater pressure concentration at the lateral aspect of the heel [26,30].  If the COP is focused to the lateral side of the calcaneus during heel strike, it is possible that the additional force required to place the individual into a compromised position may be minimal [30]. As a result, by examining the location of the COP upon initial contact it may be possible to identify running conditions that could potentially reduce the likelihood of sustaining a lateral ankle sprain by avoiding the COP displacements seen in the unstable ankles.

A commercially available design of minimalist design footwear (huaraches (HU)) have been developed (Figure 1) with minimum cushioning (4mm tread) and string uppers designed to minimally restrict natural foot movement. By comparing COP variables in participants running barefoot and wearing the HU footwear it may be possible to see the different foot mechanics in each. Therefore the aim of this study was to investigate the differences between COP variables, many of which are linked to ankle ligament injuries, measured in barefoot, huaraches and traditional foot wear runners (Figure 1). The differences in kinetics and kinematics measured between genders [19,34–37] demonstrates a need for studies investigating kinetics of locomotion to consider each gender separately and as such this research will focus on conditions during running in a healthy female population.

Methods

Selection and Description of Participants

Seventeen healthy female participants were examined (aged 21.2±2.3 years, height 165.4±5.6 cm, body mass 66.9±9.5 kg, foot size 6.8±1.0 UK). All participants were free from musculoskeletal pathology and provided written informed consent in accordance with the declaration of Helsinki.

Fig1

Figure 1 HU footwear (above) and TRS (below).

Technical Information

Participants were given time to practice running in the minimalist footwear until they felt comfortable, no prior training was undertaken [8]. Participants performed five footfalls in each footwear condition at a controlled speed of 12km/h±10% over a Footscan pressure plate (RsScan International, 1m x 0.4m, 8192 sensors) (Figure 1) collecting COP data at 250Hz positioned in the center of a 28.5m runway. Participants practiced running along the runway and adjusted their starting position to achieve a natural footstrike on the pressure mat to minimize any influence of targeting [38]. They were also instructed to look at a point on the far wall and not slow down until passing the second timing gate.

Various times (Initial Metatarsal contact (IMC), initial forefoot flat contact (IFFC, first instant all the metatarsals heads are in ground contact) and heel off (HO)) during foot to ground contact were identified (Fig.2), anterior-posterior and medial-lateral displacement and velocity data were calculated at these time points [30,39]. COP displacement and velocity values were normalized to a percentage of foot width and length as appropriate and using the same methods as in previous research [30,39]. This method of collecting COP progression data in direct foot contact and under the shoe has been confirmed as reasonable [40,41].

Statistics

Descriptive statistics including means and standard deviations were calculated for each COP variable in each condition. One way repeated measures ANOVAs were used to determine the differences between footwear conditions with significance accepted at the p<0.05 level. The Shapiro-Wilk statistic for each condition confirmed that the data were normally distributed and where the sphericity assumption was not met, correctional adjustment was made using Greenhouse-Geisser. Effect sizes were calculated using an Eta2 (η2). Post-hoc analyses were conducted using a Bonferroni correction to control type I error (Table 1). All statistical procedures were conducted using SPSS 19.0 (SPSS Inc., Chicago, IL, USA).

Results

The COP data collected was observed for each trial and various key points in time during the stance phase were identified (Figure 2)

Fig2

Figure 2 Typical BF plantar pressure.

The means were calculated for the COP timing (Table 1), COP medial-lateral (Table 2) and COP anterior-posterior (Table 3) variables.

Time variables

Analysis of the timing variables reported between the footwear conditions is displayed in Table 1 and indicated a significant main effect for the timing of IMC (F(1.41, 22.55)= 57.29, p<0.0005,  η2=0.782) and IFFC (F(2, 32)= 43.69, p<0.001,  η2=0.732) no significant effect was reported for HO (F(1.30, 20.87)= 2.56, p=0.118,  η2=0.138). Post hoc analysis revealed significant differences (p<0.001) between the TRS and both the BF and HU conditions for timing of IMC, This was also the case for the IFFC event timing which additionally reported a significant difference (p=0.04) between the BF and HU conditions.

Tab1

Table 1 Means and standard deviations of center of pressure variables timing variables. †=Significantly different (p<0.05) from BF, ¥=significantly different (p<0.05) from HU, *=significantly different (p<0.05) from TRS.

Medial Lateral COP Variables

Analysis of the movement of the COP in the Medial Lateral plane of the foot between footwear conditions are displayed in Table 2 and report that a significant main effects for the position of the COP in terms of medial lateral position (X-comp) were identified at IMC X-comp (F(1.454, 23.268= 5.87, p=0.014,  η2=0.269), IFFC X-comp (F(2, 32)= 18.9, p<0.001,  η2=0.542) and HO X-comp (F(2, 32)= 15.6, p<0.001,  η2=0.494).) No significant main effect was identified for IFCX-comp (F (2, 32) = 3.161, p=0.056, η2=0.165). Post hoc analysis revealed a significant difference for IMC X-comp (p=0.025), IFFC X-comp (p=0.001) and HO X-comp (P=0.003) between BF and TRS conditions, and a significant difference between IFFC X-comp (p<0.001) and HO X-comp (p<0.001) between HU and TRS conditions.

Significant main effects for the position of the medial lateral velocity of the COP in terms of position (VEL-X) were identified for HO VEL-X (F (2, 32) = 32.6, p<0.001, η2=0.671). Post hoc analysis revealed a significant difference for HO VEL-X between BF and TRS (p<0.001) and HU and TRS (p<0.001).  No significant main effect was identified for IMC VEL-X (F (1.46, 23.31= 1.314, p=0.279, η2=0.076) or IFFC VEL-X (F (1.33, 21.24) = 2.073, p=0.161, η2=0.115).

Tab2

Table 2 Means and standard deviations of center of medial-lateral pressure variables. †=Significantly different (P<0.05) from BF, ¥=significantly different (p<0.05) from HU, *=significantly different (p<0.05) from TRS, FW%=Percentage of foot width.

Anterior Posterior COP Variables

Analysis of the movement of the COP in the Anterior Posterior plane of the foot between footwear conditions are displayed in Table 2 and report that a significant main effects for the position of the COP in terms of anterior posterior position (Y-comp) were identified at IFCY-comp (F (2, 32) = 5.04, p<0.013, η2=0.239) and HO Y-comp (F (1.09, 17.39) = 30.71, p<0.001, η2=0.657). No significant main effect was identified for IMC Y-comp (F (1.42, 22.66) = 3.28, p=0.07,  η2=0.170) or IFFC Y-comp (F(1.22, 19.58)= 0.88, p=0.38,  η2=0.052). Post hoc analysis revealed a significant difference for HO Y-comp (p<0.001) and IFC Y-comp (p=0.025) between BF and TRS, a significant difference was also identified for HO Y-comp between HU and TRS conditions (p<0.001).

Significant main effects for the position of the anterior posterior velocity of the COP in terms of position (VEL-Y) were identified for IMC VEL-Y (F(1.41, 22.58)= 13.60, p<0.0005 η2=0.460)  and HO VEL-Y (F(1.17, 18.77)= 13.26, p=0.001,  η2=0.453) No significant main effect was identified for IFFC VEL-Y (F(1.21, 19.33)= 1.710, p=0.209,  η2=0.097). Post hoc analysis revealed a significant difference between BF and TRS for IMC VEL-Y (p=0.005) and HO VEL-Y (p=0.001), significant differences were also identified between HU and TRS for IMC VEL-Y (p=0.002) and HO VEL-Y (P=0.011).

Tab3

Table 3 Means and standard deviations of anterior-posterior center of pressure variables.†=significantly different (p<0.05) from BF, ¥=significantly different (p<0.05) from HU, *=significantly different (p<0.05) from TRS,  FL%=Percentage of foot length.

Discussion

The purpose of the current investigation was to compare the COP variables of a healthy female population running in BF, HU and TRS conditions. The first aim was to identify if there existed any differences between the shod and BF conditions, in order to identify whether running in such footwear produced similar kinetics to those found in BF running. The second aim was to determine if there were any significant differences between footwear in the COP variables implicated in the etiology of injury [30].

The significant differences in the IMC and IFFC time parameters (p<0.05) in the TRS compared to the BF and HU conditions, suggest a more plantarflexed foot placement (in BF and HU) at ground contact. This has been reported previously in analyses comparing BF to shod [2,12] and minimalist footwear compared to shod [3] conditions and suggests HU rather than TRS would be the favored footwear to reduce the incidence of injury in runners [10–12]. During running there is often uneven terrain, and as the calcaneus lands, it lends itself to movement in the coronal plane by the very nature of its shape. Furthermore, it has been identified that patients with ankle instability have a longer duration of contact from the initial heel contact to the forefoot landing [42]. Therefore, a quicker loading of the forefoot as observed in the BF and HU conditions, may offer greater support to potentially limit hazardous injury.

During locomotion, as the foot makes contact with the ground, the line of the resulting reaction force is determined by the position of the foot in relation to the athletes COG [24]. Previous research reported that when an increased angle of supination upon touchdown was present, an apparent increase in the number of ankle sprains ensued [43]. With the TRS condition in the current study exhibiting a trend towards a more laterally displaced COP, this may infer that the initial contact of the foot was made whilst being held in slight supination, and therefore similar those suffering from ankle instability which may increase susceptibility to injury.

Previous research identified that an ankle sprain group exhibited a higher loading under the medial border of the foot, and this was identified as an indicator or susceptibility to ankle sprain [30]. The significant difference between the shod and both the BF and HU condition for the IFFC X-comp variable indicated a more medially loaded foot. This may also be a predisposing factor for an inversion ankle sprain.

It appears that the HU shoe minimizes the changes in COP characteristics seen in TRS compared to BF running with only one variable (IFFC time) reporting a significant (p<0.05) difference between HU and BF. Furthermore, this particular minimalist design (HU) may more closely simulate BF running compared to some other footwear designed to simulate BF running [2]. These results suggest that proposed health benefits associated with BF running  [10] may be prevalent in HU footwear conditions.

Conclusions

The data collected in this study provides evidence that the HU design of footwear may be a suitable alternative to running BF for females, by offering protection to the plantar surface of the foot whilst adjusting the running strategy identified through COP variables in a similar way to BF running when compared to running in TRS. From a rehabilitation point of view, it may advantageous to initiate a return to running using minimalist footwear as this appears to have the potential to reduce excessive COP characteristics linked to ankle inversion injury compared to shoes. However potential injury risk reduction benefits of BF running are yet to be conclusively substantiated and any change in habitual running style through footwear choice should be approached with caution.

Future research

This study focused on a population of healthy females. Previous research has demonstrated differences between genders biomechanically and regarding injury rates [19,44] and as such the results cannot be generalized to a male sample. Therefore there is clear need to perform a similar examination using a male population. Previous research has suggested that the thickness of cushioning in running shoes may not have a significant effect on loading characteristics [7] during foot to ground impact. The HU design of shoe is commercially available in different sole thickness. Testing for similar effects of sole thickness that are observed in the HU design of shoe warrant further investigation to identify a move towards the possibility for an optimum design in the general population.

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Tibiocalcaneal kinematics during treadmill and overground running

Posted on June 30, 2014 | Comments Off on Tibiocalcaneal kinematics during treadmill and overground running

by Jonathan Sinclair1, Paul J Taylor2pdflrg

The Foot and Ankle Online Journal 7 (2): 8

Epidemiological studies analyzing the prevalence of running injuries suggest that overuse injuries are a prominent complaint for both recreational and competitive runners. Excessive coronal and transverse plane motions of the ankle and tibia are linked to the development of a number of chronic injuries. This study examined differences in tibiocalcaneal kinematics between treadmill and overground running. Ten participants ran at 4.0 m.s-1 in both treadmill and overground conditions. Tibiocalcaneal kinematics were measured using an eight-camera motion analysis system and compared using paired samples t-tests. Of the examined parameters; peak eversion, eversion velocity, tibial internal rotation and tibial internal rotation velocity were shown to be significantly greater in the treadmill condition. Therefore, it was determined treadmill runners may be at increased risk from chronic injury development.

Keywords: Biomechanics, treadmill, injury, running.

ISSN 1941-6806
doi: 10.3827/faoj.2014.0702.0008

Address correspondence to:2School of Psychology University of Central Lancashire,Preston, Lancashire, PR1 2HE.
E-mail: PJTaylor@uclan.ac.uk

1 Centre for Applied Sport and Exercise Sciences, University of Central, Lancashire.

Epidemiological studies analyzing the prevalence of running injuries suggest that overuse injuries are a prominent complaint for both recreational and competitive runners [1]. Each year approximately 19.4-79.3 % of runners will experience a pathology related to running [2].

The treadmill is now recognized as a common mode of exercise, and is becoming more popular as a running modality [3]. Since the early 1980’s the sport of running has changed dramatically, with a significant increase in the number of treadmill runners [4]. Runners’ World suggests that 40 million people in the U.S alone run using treadmills. Traditionally, treadmills have been used in clinical and laboratory research, but are now used extensively in both fitness suites and homes.

Treadmills allow ambulation at a range of velocities whilst indoors in a safe controlled environment. It is not currently known, however, whether the incidence of injuries may be affected differently between treadmill and overground running.

Lower extremity kinematic motions of excessive eversion and tibial internal rotation have been connected with various running injuries [5,6,7]. Additionally, movement coupling between the foot and shin, which causes the tibia segment to rotate internally between touchdown and midstance, has also been linked to the etiology of injury [8,9,10]. The amount of the motion transfer from ankle eversion to tibial internal rotation has been shown to differ widely among individuals [8,11]. However, given the popularity of treadmill running, surprisingly few investigations have specifically examined 3-D kinematics of the tibia and ankle during running on the treadmill in comparison to when running overground. Therefore the aim of the current investigation was to determine whether differences in tibiocalcaneal kinematics exist between treadmill and overground running.

Methods

Participants

Ten male participants (age 29.39 ± 5.17 years, height 1.81 ± 0.11m and body mass 74.19 ± 7.98kg) volunteered to take part in the current investigation. All were free from musculoskeletal pathologies at the time of data collection and provided informed consent. All runners were considered to be rearfoot strikers as they exhibited a clear first peak in their vertical ground reaction force time-curve. Ethical approval was obtained from the University Ethics Committee and the procedures outlined in the declaration of Helsinki were followed.

Procedure

All kinematic data were captured at 250 Hz via an eight-camera motion analysis system (QualisysTM Medical AB, Goteburg, Sweden). Two identical camera systems were used to collect each mode of running. Calibration of the QualisysTM system was performed before each data collection session.

The current investigation used the calibrated anatomical systems technique (CAST) [12]. To define the anatomical frame of the right; foot and shin retroreflective markers were positioned unilaterally to the calcaneus, 1st and 5th metatarsal heads, medial and lateral malleoli and medial and lateral epicondyle of the femur. A tracking cluster was positioned onto the shin segment. The foot segment was tracked using the calcaneus, 1st and 5th metatarsal markers respectively. A static trial was conducted with the participant in the anatomical position in order for the positions of the anatomical markers to be referenced in relation to the tracking markers/ clusters, following which those not required for tracking were removed.

In the overground condition participants completed ten running trials over a 22m walkway (Altrosports 6mm, Altro Ltd, Letchworth Garden City, Hertfordshire, UK) at 4.0m.s-1±5% in the laboratory. Running velocity was monitored using infra-red timing gates (SmartSpeed Ltd UK). A successful trial was defined as one within the specified velocity range, where all tracking clusters were in view of the cameras and with no evidence of gait modification due to the experimental conditions. To collect treadmill information a WoodwayTM (ELG,Weil am Rhein, Germany) high-power treadmill was used throughout. Participants were given a 5-min habitation period, in which participants ran at the determined velocity prior to the collection of kinematic data. Ten trials were also collected for treadmill kinematics. As force information was not available for each running condition, footstrike and toe-off were determined using kinematic information as in previous research [3]. The order in which participants ran in each condition was counterbalanced.

Data processing

Running data were digitized using QualisysTM Track Manager in order to identify appropriate retroreflective markers then exported as C3D files. 3-D kinematics were quantified using Visual 3-D (C-Motion Inc, Germantown, MD, USA) after marker displacement data were smoothed using a low-pass Butterworth 4th order zero-lag filter at a cut off frequency of 15 Hz [13]. 3-D kinematics were calculated using an XYZ sequence of rotations (where X represents sagittal plane; Y represents coronal plane and Z represents transverse plane rotations) [14]. All kinematic waveforms were normalized to 100% of the stance phase then processed trials were averaged. Discrete 3-D kinematic measures from the ankle and tibia which were extracted for statistical analysis were 1) angle at footstrike, 2) angle at toe-off, 3) range of motion from footstrike to toe-off during stance, 4) peak eversion/ tibial internal rotation, 5) relative range of motion (ROM) (representing the angular displacement from footstrike to peak angle, 6) peak eversion/ tibial internal rotation velocity, 7) peak inversion/ tibial external rotation velocity, 8) eversion/ tibial internal (EV/TIR) ratio which was quantified in accordance with De Leo et al [15] as the relative eversion ROM / the relative tibial internal rotation ROM.

Statistical analysis

Means and standard deviations were calculated for each running condition. Differences in the outcome 3D kinematic parameters were examined using paired samples t-tests with significance accepted at the p≤0.05 level. Effect sizes for all significant observations were calculated using a Cohen’s D statistic. The data were screened for normality using a Shapiro-Wilk test which confirmed that the normality assumption was met. All statistical analyses were conducted using SPSS 21.0 (SPSS Inc, Chicago, USA).

Results

The results indicate that while the kinematic waveforms measured during overground and treadmill running were quantitatively similar, significant differences were found to between the two running modalities. Figure 2 presents the 3-D tibiocalcaneal angular motions from the stance phase. Tables 1 and 2 present the results of the statistical analysis conducted on the tibiocalcaneal measures.

In the coronal plane, treadmill runners were associated with significantly (t (9) = 5.66, p<0.05, D= 1.22) greater peak eversion in comparison to when running overground. In the transverse plane it was also shown that peak tibial internal rotation was significantly (t (9) = 5.71, p<0.05, D= 1.28) greater when running on the treadmill compared to when running overground. Finally, the EV/ TIR ratio was shown to be significantly higher when running on the treadmill compared to overground.

TIB_table1

Table 1 Tibiocalcaneal joint angles measured during treadmill and overground running (* = significant difference).

TIB_table2

Table 2 Tibiocalcaneal angular velocities measured during treadmill and overground running (* = significant difference).

TIB1

Figure 1 Tibiocalcaneal kinematics as a function of overground and treadmill conditions (Black = treadmill and Dash = overground) (a = ankle coronal plane angle, b = tibial internal rotation angle, c = ankle coronal plane velocity, d = tibial internal rotation velocity) (EV = eversion, INT = internal).

In the coronal plane, treadmill runners were associated with significantly (t (9) = 4.65, p<0.05, D= 1.06) greater peak eversion angular velocity in comparison to when running overground. In the transverse plane it was also shown that peak tibial internal rotation angular velocity was significantly (t (9) = 4.80, p<0.05, D= 1.10) greater when running on the treadmill compared to when running overground.

Discussion

This study aimed to determine whether differences in tibiocalcaneal kinematics exist between treadmill and overground running. This represents the first comparative investigation to consider the variations that may be present in tibiocalcaneal kinematics between these two running modalities.

The key observation from the current study is that treadmill running was associated with significantly greater eversion and tibial internal parameters in comparison to overground running. This finding may relate to the deformation characteristics of the surface during the treadmill condition and has potential clinical significance. These findings suggest that running on a treadmill may be associated with an increased risk from injury as rearfoot eversion and tibial internal rotation are implicated in the etiology of a number of overuse injuries [16,17,18,19]. Therefore treadmill runners may be at a greater risk from overuse syndromes such as tibial stress syndrome, Achilles tendinitis, patellar tendonitis, patellofemoral pain, iliotibial band syndrome and plantar fasciitis [16,17,18,19].

With respect to the potential differences in coupling between ankle and tibia it was observed that treadmill running showed a trend towards having a lower ankle eversion to tibial internal rotation ratio in comparison to overground. This suggests that differences between the two running modalities may exist in terms of the distal coupling mechanism between ankle and tibia. The EV/TIR is an important mechanism as it provides insight into where an injury is most likely to occur [8]. It is hypothesized that a greater EV/TIR ratio (i.e. relatively greater rearfoot eversion in relation to tibial internal rotation) may increase the stress placed on the foot and ankle [8,20] and are thus at greater risk for foot injuries. Conversely, those with lower EV/TIR ratios (relatively more tibial motion in relation to rearfoot eversion) are at greater risk from knee related injuries [10,20,21]. As such it appears that those who habitually run on a treadmill are susceptible to knee injuries and those who train overground may be most susceptible to foot injuries.

A limitation to the current investigation was the all-male sample. Sinclair et al [22] demonstrated that females exhibited significantly greater ankle eversion compared to age matched males. Therefore future work is required to determine the influence of different running modalities in female runners. Finally, this study quantified foot kinematics using markers positioned onto the shoe may serve as a limitation of the current analysis. There is likely to be movement of the foot within the shoe itself and thus it is questionable as to whether retro-reflective markers positioned on shoe provide comparable results to those placed on the skin of the foot [23,24]. However, as cutting holes in the experimental footwear in order to attach markers to skin compromises the structural integrity of the upper [24], it was determined that the utilization of the current technique was most appropriate.

Conclusions

In conclusion, although the mechanics of treadmill and overground running have been extensively studied, the degree in which tibiocalcaneal kinematics differs between the two modalities is limited. The present study adds to the current knowledge by providing a comprehensive evaluation of tibiocalcaneal kinematics during treadmill and overground running. Given the significant increases in eversion and tibial internal rotation observed in the treadmill condition, it was determined treadmill runners may be at increased risk from chronic injury development.

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