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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

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  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.
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  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.
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  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.
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  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.
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  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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Comparison of the foot kinematics during weight bearing between normal foot feet and the flat feet

Posted on March 1, 2016 | Comments Off on Comparison of the foot kinematics during weight bearing between normal foot feet and the flat feet

by Shintarou Kudo1*, Yasuhiko Hatanaka2pdflrg

The Foot and Ankle Online Journal 9 (1): 2

Purpose: The purpose of this study was to clarify the difference in foot kinematics, including detailed analysis of the mid- and forefoot, between normal and flat feet.
Methods: Forty-six feet of 33 young normal volunteers were participated in this study. All subjects were categorized two groups which were normal foot group and flat foot group. The fifteen color markers were mounted over the anatomical landmarks. Foot motion in a stance with the measurement foot forward, the body weight loaded on the forefoot as much as possible with the lower leg inclined forward and the entire plantar surface in contact with the floor was recorded using four hi-definition digital video cameras. All markers were manually digitized using the Frame-DIAS4 software program (DKH Co. Ltd, Tokyo, Japan). The three directional coordinates of each marker were calculated, and the three directional movements of each marker were compared between normal feet and flat feet using the Man-Whitney U-test. Moreover, discriminant functional analyses were performed on all combinations if significant differences were found between normal feet and flat feet.
Results: Normal foot showed medial inclination with maintaining the arch structure, however flat foot showed that forward splaying with collapsed arch structure. Moreover, the forward movements of the cuboid and medial movement of the third metatarsal base may be key movements.
Conclusions: It might be important for the treatment of flat feet to control the stability of the lateral longitudinal arch.

Key words: flatfeet, kinematics, lateral longitudinal arch

ISSN 1941-6806
doi: 10.3827/faoj.2016.0901.0002

1 – Graduate School of Health Science, Suzuka University of medical science
Department of physical therapy, Morinomiya University of medical sciences
2 – Graduate School of Health Science, Suzuka University of medical science
Department of Physiotherapy, Suzuka University of medical science
* – Correspondence: kudo@morinomiya-u.ac.jp

During locomotion, the foot plays three important roles. The first is to buffer the impact force during the loading response, the second is to maintain stability and support the lower limb, and the third is to assist forward propulsion [1-3]. It will seem logical that changes to foot function may impair locomotion [4].

Flat feet are known to be associated with not only foot overuse problems including metatarsal stress fractures, plantar fasciitis, and Achilles tendinitis, but also knee and leg injuries such as medial tibial stress syndrome, iliotibial friction syndrome and patellofemoral pain syndrome [5-7]. It has been suggested that because the human foot is so specialized, it has limited tolerance for maladaptive disorders such as flat foot, a condition that can lead to changes in muscle functions [8]. Therefore, it is important to improve foot function by performing physical therapy such as muscle strength training of the foot intrinsic and extrinsic muscles, stretching the Achilles tendon, and the use of a foot orthosis.

Measurements of three-dimensional foot kinematics in vivo during weight bearing has been performed using a multi-segment foot model, three dimensional computed tomography. Use of a multi segment foot model enabled the measurement of motions such as ground walking, running and jumping without the need for invasive methods.  In a flat foot, analysis of three-dimensional foot kinematics during gait using the oxford foot model show that forefoot abduction movement and rear foot pronation movement are increased, while the peak plantar flexion moment increase in the late stance phase [9]. However, the forefoot is measured as one segment in this model, although the forefoot consisted of five metatarsals.

Thus, this method has a limitation in that forefoot movements can be measured in detail. Moreover, the human foot is made up of seven tarsals, five metatarsals and fourteen phalanges. These bones, together with soft tissues such as ligaments and muscles, maintain the three arches of the foot. Nester reported that the first, second and third metatarsals had greater stability compared to the fourth and fifth metatarsals, and that the fourth and fifth metatarsals were functionally distinct from the other three metatarsals [10]. Moreover, analysis of patients with flat feet show that the peak plantar flexor moment is increased during the terminal stance phase, and the peak pressure of the medial midfoot was also increased, while that of the lateral forefoot was decreased [11,12]. The terminal stance phase is shown the raising the heel and load on only the forefoot. Therefore, it is necessary to investigate the forefoot kinematics of the flat foot in detail during forefoot loading. The purpose of this study is to clarify the characteristic of the foot kinematics including detailed analysis of the mid- and forefoot during forefoot loading in the flat foot. We hypothesis that both metatarsal and tarsals movement of the flat foot are larger than those of the normal feet.

Material and methods

Forty-six feet of 33 young normal volunteers (seventeen males and sixteen females 22.0 ± 4.0 years old) who had provided informed consent were involved in this study. All subjects were categorized into one of two groups: a normal foot group and a flat foot group, using both the Foot Posture Index-6 and a medical history of pain related to flat feet. 26 feet of 21 persons (10 males and 11 Females, 21.0 ± 2.4 years old) constituted the normal foot group in which the height of the medial longitudinal arch was maintained in the standing position without any pain history related to flat foot, and 20 feet of 12 persons (7 males and 5 Females, 23.8 ± 5.6 years old) constituted the flat foot group in which the height of the medial longitudinal arch had collapsed with some pain history related to the flat foot. All subjects were free from lower limb injuries and pain at the time of testing. The 16 flat feet without any pain history and 4 normal feet with some foot pain history are excluded in this study. This study was approved by the ethics committee of our university (2014-078).

Fifteen skin color markers were attached over the following anatomical bony landmarks; the five metatarsal heads (MTH1~5) and bases (MTB1~5), the navicular (nav), the cuboid (cub), the peroneal trochlea (cal-l), the sustentaculum tali (cal-m), and the posterior tip of the calcaneus (cal-p). The subjects performed the forward lunge without a stride. Starting position of the forward lunge involved standing upright with measurement foot stance one step forward onto a sheet censer (Win-pod, Medicapteurs s.a.s., France) which could measure the plantar pressure distribution and set on a force plate(Anima co., Japan). The measurement side of the knee and the ankle slowly flexed until both maximum dorsi-flexion of the ankle and forefoot weight bearing (Figure 1). Subjects were instructed the time of the forward lunge was almost 1second, and enough practice was performed. Forward lunge was performed 3 repetitions. Both whole plantar surface in contact with floor and approximate 70-80 percent of the body weight was loaded on the forefoot were confirmed using the plantar pressure distribution with 180Hz. And the vector of the ground reaction force was confirmed using the force plate with 180Hz.

Foot motions during the forward lunge were recorded using four hi-definition digital video cameras (GZ-G5. Victor Co., Tokyo, Japan) with 60Hz. The global coordination frame was set using an acrylic cube with sides of 64.6 millimeters in length, and the direction of movement was represented on the y axis, the position in the vertical direction was represented on the z axis, and the axis at right angles between the y axis and the z axis was represented on the x axis.

F1

Figure 1 Forward lunge without stride. The measurement side (right side) of the knee and the ankle slowly flexed until both maximum dorsi-flexion of the ankle and forefoot weight bearing. Both whole plantar surface in contact with floor and approximate 70-80 percent of the body weight was loaded on the forefoot are confirmed using both the plantar pressure distribution and the force plate.

All markers were manually digitized and three dimensional markers reconstructed using the Frame-DIAS4 software program (DKH Co. Ltd, Tokyo, Japan). Three dimensional coordinates of each marker were calculated, and low-pass filtered using a Butterworth digital filter with a cut-off frequency of 10 Hz. This method had high accuracy (root mean square error was 0.39 mm) in a pilot study. The three directional coordinates of each marker were calculated, and the three directional movements of each marker were compared between normal feet and flat feet using the Man-Whitney U-test. Moreover, discriminant functional analyses were performed on all combinations if significant differences were found between normal feet and flat feet using SPSS statistics software ver.18 (IBM, Armonk, NY, USA).

F2

Figure2 Scatter diagram of the discriminant functional analysis. White rhombus; normal foot, Black circle; the flat foot. Both the MTB3(x) and the Cub(y) were the most powerful variables in distinguishing between normal foot and flat feet foot. Based on these two variables 90.5% of the patients were correctly classified (Wilk’s lambda 0.3; chi-squared 37.7; P < 0.01). If the step height was additionally included in the discriminant analysis, 70% of the flat feet foot were correctly classified.

Results

In the medial direction, flat feet are significantly smaller than normal feet at the first, third, and fourth metatarsal head, the first to third metatarsal bases, the navicular and the cuboid. In the forward direction, flat feet are significantly larger than normal feet at the first metatarsal head and base, the second and third metatarsal base, the cuboid and the cal-p. In the vertical direction, flat feet are significantly smaller than normal feet at the first, fourth and fifth metatarsal head, the fifth metatarsal base, the navicular and the cuboid, and larger at the second metatarsal head and cal-p (Table 1). In discriminant analysis, between the groups the cub(y) and MTB3(x) were the most powerful variables in distinguishing between normal feet and flat feet. Based on these two variables 90.5% of the patients were correctly classified (Wilk’s lambda 0.3; kai square 37.7; P < 0.01).

 

table1

Table1 Differences between the normal foot feet and the flat foot feet in each direction. Median value and 25, 75 percentile value were shown. All measurements were in millimeters, *; statistic significant statistically significant differences. [MTH1; First metatarsal head, MTH2; Second metatarsal head, MTH3; Third metatarsal head, MTH4; Fourth metatarsal head, MTH5; Fifth metatarsal head, MTB1; First metatarsal base, MTB2; Second metatarsal base, MTB3; Third metatarsal base, MTB4; fourth metatarsal base, MTB5; fifth metatarsal base, Nav; Navicular, Cub; cuboid, Cal-m; the sustentaculum tali, Cal-l; the peroneal trochlea, Cal-p; calcaneus posterior.]

If the step height was additionally included in the discriminant analysis, 70% of flat feet were correctly classified (Figure 2).

Discussion

Our results are shown the characteristic of the flat feet and normal feet during the forefoot loading. We hypothesized that both metatarsals and tarsals movements of the flat foot were larger than those of the normal feet. However, some makers of the flat foot are larger than normal foot in the forward direction, and are smaller than those in normal foot in the medial-lateral direction in this study. Therefore, we consider that there are difference characteristics of the foot kinematics during the forefoot loading between normal foot and flatfoot. Previous studies of foot movement during weight bearing were divided into three mechanisms which were the “truss mechanism” [13], the “twisted foot plate model” [14], and the “mid tarsal locking mechanism” [15]. These models and mechanisms had shown that the rear foot was pronated, and the medial longitudinal arch of the mid- and forefoot was in dorsal flexion, when the body weight was loaded onto the foot. However, the truss mechanism was based on the two dimensional analysis. And in the twisting foot plate model and the mid tarsal locking mechanism, forefoot were represented one segment, while forefoot was consisted with five metatarsals. And these models did not discuss differences in foot kinematics between normal feet and flat feet, and it was considered that flat feet were a result of excessive pronation.

Our results of normal feet are shown that all makers exclude the cal-p drop in antero-medially, and the cal-p moved to superior. Moreover, each metatarsal base makers move larger than each metatarsal head maker, respectively. Thus, we find that the calcaneus is inclined medially with plantar flexion and the whole foot drop and incline medially with forefoot dorsiflexion during forefoot loading. Thus, we call the normal feet movement during the forefoot loading “medial inclination”. Sarrafian reported that the foot was similar to a twisted plate. The posterior segment of the plate was compressed side-to-side whereas the anterior segment was compressed in a dorso-plantar direction, and the twisting of the plate determines a longitudinal and a mid-segment transverse arch [14]. Thus, the foot movements we find are nearly equal “twisted foot plate model”.

In the flat foot, forward movements of the makers are larger than normal foot with whole foot medial inclination. Decreasing the function of the arch support structure led to collapse medial longitudinal arch without maintenance the rigidity of the foot. Thus, we call this movement with collapse of the medial longitudinal arch structure during the forefoot loading “forward splaying”.

A traditional foot orthosis was inserted to limit calcaneal pronation, and to support the medial longitudinal arch using a heel wedge and a navicular pad. These foot orthoses focused on control of excessive foot pronation. However, our results are shown that both the forward movements of the cuboid and medial movement of MTB3 are key movements to control foot kinematics for the flat foot. Lundgren reported that the mobility of the lateral side of the foot (i.e. the cuboid and the fifth metatarsal) was greater than that of bones of the medial side of the forefoot such as the medial cuneiform and the first metatarsal [16]. This indicated that it might be important for the treatment of flat feet to control the stability of the lateral longitudinal arch. In particular, the treatment of flat feet aimed to decrease forward movement of the cuboid, and to increase medial movement of the third metatarsal base may have possibility to improve the foot function for flat feet deformity.

This study had a limitation. The subjects of the flat foot group did not have any symptom. It has possible that symptomatic subject show the abnormal movement which is caused by the some pain of the foot, and we cannot find the abnormal foot motions relate to the flat foot. Therefore, we observe the subject without symptom, however patients with both flat foot and some symptom may show the difference foot movement with this study.

References

[1] Whittle M. Generation and attenuation of transient impulsive forces beneath the foot: a review. Gait Posture 1999; 10: 264-275. (PubMed)
[2] Winter D. Human balance and posture control during standing and walking. Gait Posture 1995; 3: 193-214.(http://www.sciencedirect.com/science/article/pii/0966636296828499)
[3] Stefanyshyn DJ, Nigg BM. Mechanical energy contribution of the metatarsophalangeal joint to running and sprinting. J Biomech 1997; 30: 1081-5.(PubMed)
[4] Arnold JB, Mackintosh S, Jones S, Thewlis D. Differences in foot kinematics between young and older adults during walking. Gait Posture 2014; 39(2): 689-94. (PubMed)
[5] Kaufman KR, Brodine SK, Shaffer RA, Johnson CW, Cullison TR. The effect of foot structure and range of motion on musculoskeletal overuse injuries. Am J Sports Med 1999; 27(5): 585-93. (PubMed)
[6] Simkin A, Leichter I, Giladi M, Stein M, Milgrom C. Combined effect of foot arch structure and an orthotic device on stress fractures. Foot Ankle 1989; 10(1): 25-9. (PubMed)
[7] Heil B. lower limb biomechanics related to running injuries. Physiotherapy 1992; 78(6): 400-406.  (http://www.sciencedirect.com/science/article/pii/S0031940610615246)
[8] Olson TR, Seidel MR. The evolutionary basis of some clinical disorders of the human foot: a comparative survey of the living primates. Foot Ankle 1983; 3(6): 322-41. (PubMed)
[9] Levinger P, Murley GS, Barton CJ, Cotchett MP, McSweeney SR, Menz HB. A comparison of foot kinematics in people with normal- and flat-arched feet using the Oxford Foot Model. Gait Posture 2010; 32(4): 519-23. (PubMed)
[10] Nester CJ. Lessons from dynamic cadaver and invasive bone pin studies: do we know how the foot really moves during gait? J Foot Ankle Res 2009; 2:18. (PubMed)
[11] Hunt AE, Smith RM. Mechanics and control of the flat versus normal foot during the stance phase of walking. Clin Biomech 2004; 19(4): 391-7. (PubMed)
[12] Chuckpaiwong B, Nunley JA, Mall NA, Queen RM. The effect of foot type on in-shoe plantar pressure during walking and running. Gait Posture 2008; 28(3):405-11. (PubMed)
[13] Hicks JH. The mechanics of the foot: the plantar aponeurosis and the arch. J Anat 1954; 88:25-30. (PubMed)
[14] Sarrafian SK. Functional characteristics of the foot and plantar aponeurosis under tibiotalar loading. Foot & Ankle 1987; 8(1): 4-18. (PubMed)
[15] Blackwood CB, Yuen TJ, Sangeorzan BJ, Ledoux WR. The midtarsal joint locking mechanism. Foot Ankle Int 2005; 26(12): 1074-80. (http://fai.sagepub.com/content/26/12/1074.short)
[16] Lundgren P, Nester CJ, Liu A, Arndt A, Jones R, Stacoff A, et al. Invasive in vivo measurement of rear- mid- and forefoot motion during walking. Gait Posture 2008; 28(1): 93-100. (PubMed)

 

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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.

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Differences in multi-segment foot kinematics measured using skin and shoe mounted markers

Posted on June 30, 2014 | Comments Off on Differences in multi-segment foot kinematics measured using skin and shoe mounted markers

by Jonathan Sinclair1, PJ Taylor2, J Hebron3, N Chockalingam4pdflrg

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

Models with three segments have been implemented in order to represent the movement of the foot in a comprehensive way during walking and running, however the efficacy of mounting such a system of markers externally onto the shoe has not been explored. The aim of the current investigation was to determine whether 3-D three-segment foot kinematics differ between skin and shoe-mounted markers. Twelve male participants walked and ran at 1.25m/s and 4.0m/s along a 22 m runway. Multi-segment foot kinematics were captured simultaneously using markers placed externally on the shoe and on the skin through windows cut in the shoe. Wilcoxon tests were used to compare the 3-D kinematic parameters, and coefficients of multiple correlations (CMC) were employed to contrast the 3-D kinematic waveforms. Strong correlations were observed between the calcaneus-tibia waveforms R2 ≥0.957. However, at the more distal foot articulations lower correlations were found midfoot-calcaneus R2 ≥0.484, metatarsus-midfoot R2 ≥0.538 and metatarsus-calcaneus R2 ≥0.335. Significant differences between in discrete kinematic parameters were also observed between skin and shoe mounted markers, at the midfoot-calcaneus, forefoot-midfoot and forefoot-calcaneus articulations. The results indicate that shoe mounted markers do not fully represent true foot movement, and should therefore be interpreted with caution during examination of multiple-segment foot kinematics.

Keywords: Multi-segment foot, biomechanics, kinematics, overuse injury.

ISSN 1941-6806
doi: 10.3827/faoj.2014.0701.0001

Address correspondence to: 1Jonathan Sinclair,
Division of Sport, Exercise and Nutritional Sciences, University of Central Lancashire, Preston, Lancashire, PR1 2HE.
E-mail: JKSinclair@uclan.ac.uk

2 School of Psychology, University of Central Lancashire, Preston, Lancashire, PR1 2HE. E-mail: PJTaylor@uclan.ac.uk
3 Division of Sport Exercise and Nutritional Sciences, University of Central Lancashire
4 Faculty of Health Sciences, Staffordshire University

During three-dimensional (3-D) kinematic analyses of gait biomechanical models traditionally represent the foot as a single rigid segment [1]. However, more recently three-segment foot models have been implemented in order to represent the movement of the foot in a detailed manner during walking and running analyses [2].

To quantify foot movements, retro-reflective markers were attached either to the skin or through external palpation to the shoe surface [3,4]. The accuracy of both techniques has been shown to be acceptable in clinical situations, with the majority of errors being <5mm [5]. The efficacy of the shoe mounted technique has been questioned during analyses using both single and multi-segment foot models [3,4]. During dynamic movements, such as walking and running, the foot may move inside the shoe resulting in larger inaccuracies in actual foot position measurements [3,4,6]. Therefore, inaccuracies typically referred to as movement artefact, may be introduced as a function of this movement [1].

Several procedures have been established in an attempt to overcome the potential inaccuracies associated with placing markers on the shoe. Markers attached directly to the underlying bone structures using Kirschner bone pins are utilized to accurately quantify underlying skeletal movement [7]. This technique is extremely limited due to its invasiveness and concerns regarding the ecological validity of gait patterns following the attachment of surgical equipment under local anaesthetic. The most utilized non-invasive technique is to position markers onto the foot through windows cut into the experimental footwear [6].

Previous investigations have examined the 3-D kinematic differences between shoe and skin mounted markers when using a single segment foot model. Sinclair et al examined the differences in stance phase kinematics between markers positioned onto the shoe and those positioned inside windows cut into the shoe [1]. The study documented that eversion range of motion, peak eversion, peak transverse plane range of motion, velocity of external rotation and peak eversion velocity were all significantly underestimated using shoe-mounted markers. However, there is clear paucity of studies that have examined these differences when using more complex multi-segment foot models. The aim of the current investigation was to compare the 3-D three-segment foot kinematics between skin- and shoe-mounted markers. This study tests the hypothesis that significant differences between skin and shoe mounted markers will be observed.

Methods

Participants

Twelve healthy male participants (age 24.23 SD 2.22 y, height 1.74 m SD 0.10, mass 75.78 SD 6.90 kg) were recruited for 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

Kinematic parameters were obtained at 250 Hz using an eight-camera motion analysis system (Qualisys Medical, Sweden) whilst participants walked and ran at 1.25m/s and 4.0m/s along a 22 m runway. Participants struck a Kistler 9281CA (Kistler Instruments, UK) embedded force platform [8] sampling at 1000 Hz with their dominant foot in order to define gait events of footstrike and toe-off. The stance phase was determined as the time over which a 20 N or greater vertical force was applied to the force platform [9].

Markers were placed on anatomical landmarks in accordance with the Leardini et al [2] foot model protocol allowing the anatomical frames of the calcaneus, midfoot and forefoot to be defined. The calibrated anatomical systems technique (CAST) procedure for modelling and tracking segments was adhered to [10]. Windows were cut in the laboratory-supplied experimental footwear (Pro Grid Guide 2, Saucony, USA) at the approximate locations of those outlined by Leardini et al [2]. The pre-established guidelines for length and width outlined by Shultz & Jenkyn were adhered to [11]. The three foot segments were simultaneously tracked using markers positioned on the shoe and also those on the skin within the shoe windows. Additional markers were positioned on the medial and lateral femoral epicondyles to allow the anatomical frame of the tibia to be delineated and a rigid tracking cluster was also positioned on the tibia.

Data processing

Data were digitized using Qualisys track manager and exported to Visual 3D (C-motion, Germantown USA). Marker trajectory data were filtered at 6 and 12 Hz for walking and running trials [12]. Stance phase joint angles were computed using and XYZ sequence of rotations between the calcaneus-tibia (Cal-Tib), midfoot-calcaneus (Mid-Cal), forefoot-midfoot (Fore-Mid) and forefoot-calcaneus (Fore-Cal).

Statistical analysis

Descriptive statistics for the stance phase peak angles (PK) and range of motion (ROM) for both skin and shoe mounted markers were computed, including the mean differences between the two techniques. The similarity of stance phase waveforms was examined using coefficient of multiple correlations (CMC) in accordance with the procedure outlined by Ferrari et al [13]. Based on predominant non-normality of the dataset differences in stance phase kinematic parameters were examined using Wilcoxon rank tests with the alpha criterion for statistical significance adjusted to p=0.002 based on the number of comparisons to control type I error. Statistical procedures were undertaken using SPSS v21 (IBS, SPSS INC USA).

 

MSEG1

Figure 1 Multi-segment foot kinematics during running in the a. sagittal, b. coronal and c. transverse planes as a function of skin and shoe mounted markers (Black = shoe and Dot = skin).

Results

The results indicate that the 3-D kinematic curves measured using the shoe and skin-mounted markers were in the main quantitatively similar, although significant differences were found to exist in discrete kinematic parameters. Figures 1-2 present the 3-D angular motions of the multi-segment foot during the stance phase of both running and walking. Table 1 presents the results of the statistical analysis conducted on the joint angle measures and Table 2 shows the similarity between skin and shoe mounted waveforms measured using CMC.

Discussion

The aim of the current investigation was to compare the 3-D three-segment foot kinematics between skin and shoe-mounted markers. This study represents the first to statistically examine the differences in stance phase waveforms and discrete kinematic parameters. The 3-D kinematics of the foot segments during walking and running are of great interest in both biomechanical and clinical examinations of patients [1,14,15]. Kinematic marker sets are commonly used to quantify the foot and ankle mechanics during gait and have interchangeably been applied to both the skin surface of the foot and on the shoe surface with little consideration for accuracy in the latter condition [1].

MSEG2

Figure 2 Multi-segment foot kinematics during walking in the a. sagittal, b. coronal and c. transverse planes as a function of skin and shoe mounted markers (Black = shoe and Dot = skin).

In support of the hypothesis, the results of the current investigation show that significant differences in discrete three-segment foot kinematic parameters were observed between shoe and skin mounted markers during both running and walking. It is important to note that there were significant differences between the two marker configurations in all three planes of rotation. These differences were observed primarily at the more distal articulations with the largest deviations being noted at the Fore-Mid complex. Notably, the findings of the current study oppose the single segment foot investigation of Sinclair et al, who showed that the shoe mounted markers served to underestimate foot movements, whilst in the current investigation there was a trend towards overestimation [1]. It was hypothesized that that this divergence may relate to the errors in experimental kinematic data due to violation of the rigid body in single segment foot analyses, which would be proliferated when quantifying multi-segment foot kinematics.

The greatest similarity between 3-D kinematic curves was demonstrated at the Cal-Tib complex. However, for the relative Fore-Cal and Mid-Cal rotations there was generally a low level of similarity between the two tracking techniques. It was hypothesized that these observations may relate to the poorer fit of footwear that has been observed in the more distal aspects of the foot due to its natural curvature [16]. As the fit is poorer in these regions the relative foot-shoe movement is likely to be larger thus resulting in a lack of agreement when these regions of the foot are quantified simultaneously using shoe and skin mounted markers.

MSEG_table1

Table 1 Multi-segment foot kinematics obtained as a function of skin and shoe mounted markers.

MSEG_table2

Table 2 Coefficient of multiple correlations for 3-D joint waveforms.

The current investigation also shows that during running there was lower similarity between skin and shoe mounted markers. The mean differences in discrete kinematic parameters between shoe and skin mounted markers were also larger during running than when walking. It is likely that this relates to the increased relative motion in all three planes of rotation during running in comparison to walking [17]. The increased motion of the foot segments relative to one another during running is likely to increase the propensity for relative foot-shoe movement, decreasing the similarity between foot kinematics quantified using markers placed on the shoe and those positioned onto the foot itself.

The current study further substantiates the notion that markers positioned on the shoe do not represent true foot movement when contrasted against markers placed onto the skin. The observations from this study may have clinical significance as malalignment and dysfunction of the foot articulations have been associated with an increased incidence of overuse and traumatic injury in athletes [18,19,20]. As such, misrepresentation may serve to confound the efficacy of epidemiological analyses.

Conclusions

Although previous studies have compared shoe to skin-mounted markers, current knowledge is still limited in terms of the parameters that have been taken under consideration. This study adds to the literature by providing a comprehensive 3-D kinematic and waveform comparison between skin and shoe-mounted foot models. Given that significant differences were observed between skin and shoe-mounted markers in key coronal and transverse plane parameters, it can be concluded that the results of studies using shoe-mounted markers should be interpreted with caution, particularly when performing clinical analyses. Future analyses may consider placing markers onto the skin surface through appropriately sized holes in experimental footwear.

References

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