Tag Archives: biomechanics

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.

Effects of high and low cut on Achilles tendon kinetics during basketball specific movements

by Jonathan Sinclair1*, Benjamin Sant1pdflrg

The Foot and Ankle Online Journal 9 (4): 5

The aim of the current investigation was to examine the influence of high and low-cut specific basketball footwear in relation to minimalist and conventional athletic footwear on the loads experienced by the Achilles tendon during basketball specific movements. Ten males performed run and 45˚ cut movements whilst wearing low-cut, high-cut, minimalist and conventional athletic footwear. Achilles tendon forces were calculated using Opensim software allowing the magnitudinal and temporal aspects of the Achilles tendon force to be quantified.  Differences in Achilles tendon load parameters were examined using 4 (footwear) x 2 (movement) repeated measures ANOVA. The results show that a main effect was evident for peak Achilles tendon force, which was significantly larger in the minimalist (run = 5.74 & cut = 5.85 BW) and high-cut (run = 6.63 & cut = 6.01 BW) footwear in relation to the low-cut (run = 5.79 & cut = 5.47 BW) and conventional (run = 5.66 & cut = 5.34 BW) conditions. In addition a main effect was also evident for Achilles tendon load rate, which was significantly larger in the minimalist (run = 48.84 & cut = 43.98 BW/s) and high-cut (run = 54.31 & cut = 46.51 BW/s) footwear in relation to the low-cut (run = 43.15 & cut = 31.57 BW/s) and conventional (run = 44.74 & cut = 31.15 BW/s) conditions. The current investigation indicates that minimalist and high-cut footwear may place basketballers at increased risk for Achilles tendon pathology as a function of their training/ competition. Furthermore, it appears that for basketballers who may be susceptible to Achilles tendinopathy that low-cut and conventional conditions are most appropriate.

Keywords: basketball, Achilles tendon, biomechanics

ISSN 1941-6806
doi: 10.3827/faoj.2016.0904.0005

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


At all levels of play basketball is becoming a uniquely popular athletic discipline throughout the world [1]. Basketball is regarded as a physiologically demanding sport in which players are required to perform a series of different motions that typically include running, jumping and rapid changes of direction [2]. A typical competitive basketball season will require players to train frequently and perform >60 games, a regimen which serves to place high physical and mechanical demands on those involved [3].

Basketball has in recent years gained more research attention from the scientific community regarding players’ susceptibility to injury. Research investigating the prevalence of injuries in basketball players has shown that in relation to other non-contact sports basketball is associated with a comparatively high rate of injury. Information from aetiological analyses indicates that 11.6 injuries occur per 1000 appearances, and that the vast majority (65 %) are confined to the lower extremities [4]. Athletic disciplines which include frequently jumps, foot strikes and changes in direction such as basketball, place high loads on the Achilles tendon placing it at high risk from injury [5].

Given the highly physical nature of modern basketball, court footwear must now fulfill a range of biomechanical parameters such as traction, support, stability and shock attenuation [6]. Traditionally basketball specific footwear designs were available only with high-cut ankle supports which are utilized in order to promote mediolateral stability during landing [7]. In recent times however, low-cut footwear models have also been introduced and utilized at all levels of play, meaning court specific footwear can be selected based on individual preference. Recreational level players are also known to use low-cut conventional athletic footwear which may serve to enhance improve impact loading but at the expense of medio-lateral stability [7]. In comparison to other sports such as running there is currently a paucity of scientific research examining the efficacy of basketball footwear.

Appropriate footwear selection has been cited as a mechanism by which the risk from Achilles tendon pathologies during sport can be mediated. Considerable research has examined the effects of different footwear on the forces experienced by the tendon during different sports. Sinclair examined the effects barefoot and in minimalist footwear on Achilles tendon kinetics in relation to conventional running shoes [8]. Their results showed that conventional footwear significantly reduced peak Achilles tendon forces in relation to barefoot and minimalist conditions. Similarly, Sinclair et al., [9] examined the effects of minimalist and netball specific conditions on the forces experienced by the Achilles tendon during running and cutting movements. They showed that the peak force and rate of force application was significantly reduced in the netball specific condition. Finally, Sinclair et al [10] investigated the effects of minimalist energy return and convention athletic footwear on Achilles tendon loads during depth jumping. They showed the footwear did not significantly affect Achilles tendon forces during this movement. However, despite the wealth of peer reviewed literature examining the effects of different footwear on Achilles tendon kinetics there is currently no information available regarding the influence of basketball specific shoes.

Therefore, the aim of the current investigation was to examine the influence of high and low-cut specific basketball footwear in relation to minimalist and conventional athletic footwear on the loads experienced by the Achilles tendon during basketball specific movements. The findings from the current investigation may provide basketball players with important clinical information regarding the selection of appropriate footwear, which may ultimately help to attenuate their risk from developing Achilles tendon pathologies.

Methods

Participants

Ten male participants, volunteered to take part in this study. 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 24.26 ± 4.05 years, height 1.77 ± 0.07 cm and body mass 78.66 ± 7.43 kg. The procedure utilized for this investigation was approved by the University of Central Lancashire, Science, Technology, Engineering and Mathematics, ethical committee.

Footwear

The footwear used during this study consisted of minimalist (Vibram five-fingers Original;), high-cut (Nike Lebron XII), low-cut (Nike Lebron XII Low) footwear and conventional (New Balance 1260 v2) (shoe size 9–10 in UK men’s sizes).

Procedure

Participants completed five repeats of two sport specific movements; run and cut in each of the four footwear conditions. To control for any order effects the order in which participants performed in each footwear/ movement condition were counterbalanced. Kinematic information from the lower extremity joints was obtained using an eight camera motion capture system (Qualisys Medical AB, Goteburg, Sweden) using a capture frequency of 250 Hz. To measure kinetic information an embedded piezoelectric force platform (Kistler National Instruments, Model 9281CA) operating at 1000 Hz was utilized. The kinetic and kinematic information were synchronously obtained and interfaced using Qualisys track manager.

To define the anatomical frames of the thorax, pelvis, thighs, shanks and feet retroreflective markers were placed at the C7, T12 and xiphoid process landmarks and also positioned bilaterally onto the acromion process, iliac crest, anterior superior iliac spine, posterior superior iliac spine, medial and lateral malleoli, medial and lateral femoral epicondyles, greater trochanter,  calcaneus, first metatarsal and fifth metatarsal. Carbon-fibre tracking clusters comprising of four nonlinear retroreflective markers were positioned onto the thigh and shank segments. 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. A static trial was conducted with the participant in the anatomical position in order for the anatomical positions to be referenced in relation to the tracking markers, following which those not required for dynamic data were removed.

Data were collected during the run and cut movements according to below procedures:

Run

Participants ran at 4.0 m.s-1 ±5% and struck the force platform with their right (dominant) limb. The average velocity of running was monitored using infrared timing gates (SmartSpeed Ltd UK). The stance phase of running was defined as the duration over > 20 N of vertical force was applied to the force platform[11].

Cut

Participants completed 45° sideways cut movements using an approach velocity of 4.0 m.s-1 ±5% striking the force platform with their right (dominant) limb. In accordance with McLean et al.,[12] cut angles were measured from the centre of the force plate and the corresponding line of movement was delineated using masking tape so that it was clearly evident to participants. The stance phase of the cut-movement was similarly defined as the duration over > 20 N of vertical force was applied to the force platform [11].

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 kinematic data were smoothed using cut-off frequencies of 25 and 12 Hz with a low-pass Butterworth 4th order zero lag filter.

Data during the stance phase were exported from Visual 3D into OpenSim software (Simtk.org), which was used give to simulations of muscles forces. Simulations of muscle forces were obtained using the standard gait 2392 model within Opensim v3.2. This model corresponds to the eight segments that were exported from Visual 3D and features 19 total degrees of freedom and 92 muscle-tendon actuators.

We firstly performed a residual reduction algorithm (RRA) within OpenSim, this utilizes the inverse kinematics and ground reaction forces that were exported from Visual 3D. The RRA calculates the joint torques required to re-create the dynamic motion. The RRA calculations produced route mean squared errors <2°, which correspond with the recommendations for good quality data.  Following the RRA, the computed muscle control (CMC) procedure was then employed to estimate a set of muscle force patterns allowing the model to replicate the required kinematics 13. The CMC procedure works by estimating the required muscle forces to produce the net joint torques.

Achilles tendon force was estimated in accordance with the protocol of Almonroeder et al [14] by summing the muscle forces of the medial gastrocnemius, lateral, gastrocnemius, and soleus muscles. Achilles tendon load rate was quantified as the peak Achilles tendon force divided by the time to peak force. All Achilles tendon load parameters were normalized by dividing the net values by body weight (BW).

Analyses

Differences in kinetic and kinematic parameters between footwear were examined using 4 (footwear) x 2 (movement) repeated measures ANOVAs, with significance accepted at the P≤0.05 level. Effect sizes were calculated using partial eta2 (pη2). Follow up comparisons on significant interactions were examined using simple main effects and 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

Tables 1 and Figure 1 present the footwear differences in Achilles tendon kinetics both movements. The results indicate that the experimental footwear significantly affected Achilles tendon load parameters.

 

Minimalist High-cut Low-cut Conventional
Run Cut Run Cut Run Cut Run Cut
Mean SD Mean SD Mean SD Mean SD Mean SD Mean SD Mean SD Mean
Peak Achilles tendon force (BW) 5.74 0.75 5.85 1.03 6.63 1.19 6.01 0.69 5.79 0.78 5.47 1.00 5.66 0.90 5.34
Time to peak Achilles tendon force (s) 0.12 0.01 0.16 0.04 0.13 0.02 0.16 0.02 0.14 0.01 0.19 0.04 0.13 0.02 0.18
Achilles tendon load rate (BW/s) 49.84 8.70 43.98 18.68 54.31 17.49 46.51 14.71 43.15 9.16 31.57 11.96 44.74 11.97 31.15

Table 1 Achilles tendon kinetics as a function of footwear and movement conditions.

fig1

Figure 1 Achilles tendon kinetics during the stance phase (a. = run & b. = cut) (black = minimalist, black dash = high-cut, grey dot = low-cut & grey = conventional).

For peak Achilles tendon force a significant main effect (P<0.05, pη2 = 0.64) was observed for footwear. Post-hoc pairwise comparisons showed that peak Achilles tendon force was significantly larger in the high-cut footwear in relation to the minimalist, low-cut and conventional athletic conditions. In addition it was also revealed that peak force was significantly larger in the minimalist footwear in comparison to the conventional condition.

For time to peak Achilles tendon force significant main effects were observed for both footwear (P<0.05, pη2 = 0.55) and movement (P<0.05, pη2 = 0.70). Post-hoc analysis for footwear showed that time to peak force was significantly greater in the low-cut footwear in comparison to the minimalist, high-cut and conventional conditions. Furthermore, it was also demonstrated that time to peak force was significantly greater in the conventional athletic footwear in relation to the minimalist and high-cut conditions. Finally, it was shown that time to peak force was significantly greater in the high-cut footwear in comparison to the minimalist condition. In addition post-hoc analysis for movement indicated that time to peak Achilles tendon force was significantly greater when performing the cut movement.

For Achilles tendon load rate significant main effects were observed for both footwear (P<0.05, pη2 = 0.42) and movement (P<0.05, pη2 = 0.47). Post-hoc analysis for footwear showed that Achilles tendon load rate was significantly larger in the minimalist and high-cut footwear in relation to the low-cut and conventional conditions.  

Discussion

The current study aimed to examine the effects of different basketball footwear on the loads experienced by the Achilles tendon during sport specific movements. To the authors knowledge this investigation is the first comparative examination of the effects of different footwear on Achilles tendon kinetics during basketball specific movement. The findings from this work may provide basketball players with important information regarding the selection of appropriate footwear to attenuate their risk from developing Achilles tendon pathologies.

The primary observation from the current work is that Achilles tendon loading parameters were shown to be significantly larger in the minimalist and high-cut footwear in comparison to the conventional low-cut conditions. This observation is in agreement with those of Sinclair [8] and Sinclair et al [9] who showed that minimalist footwear were associated with significant increases in Achilles tendon loading.

This observation may provide important clinically meaningful information regarding the aetiology of Achilles tendon pathologies. Achilles tendon pathologies are considered to be initiated by high loads which are experienced too frequently by the tendon itself 15. Tendon loading at an appropriate level can initiate collagen synthesis and positively influence the mechanical properties of the tendon [16]. However, when mechanical loads exceed the physiological threshold for collagen synthesis and the remodeling threshold is exceeded, this facilitates tendon degradation and ultimately leads to injury [16]. Therefore the findings from the current investigation indicate that minimalist and high-cut footwear may place basketballers at a greater risk from Achilles tendon pathologies as a function of their training/ competition.

In conclusion, although the effects of different footwear on Achilles tendon forces have been examined previously, our current knowledge of differences in Achilles tendon kinetics when performing sport specific movements in basketball footwear is limited. The current study therefore sought to provide an evaluation of Achilles tendon forces when performing sport specific movements in different basketball specific footwear. This work shows importantly that peak Achilles tendon force and the rate of Achilles tendon load rate were significantly larger in minimalist and high-cut footwear in relation to the low-cut and conventional conditions. As such given the association between Achilles tendon loading and tendon pathology the current investigation indicates that minimalist and high-cut footwear may place basketballers at increased risk for Achilles tendon pathology as a function of their training/ competition. Furthermore, it appears that for basketballers who may be susceptible to Achilles tendinopathy that low-cut and conventional conditions are most appropriate.

References

  1. Cumps, E, Verhagen R, and Meeusen R. “Prospective epidemiological study of basketball injuries during one competitive season: ankle sprains and overuse knee injuries.” J Sport Sci Med 6: 204-211, 2007. https://www.ncbi.nlm.nih.gov/pmc/articles/PMC3786241/
  2. Montgomery PG, Pyne DB and Minahan CL. The physical and physiological demands of basketball training and competition. Int J Sports Physiol Perform. 15: 75-86, 2010. https://www.ncbi.nlm.nih.gov/pubmed/20308698
  3. Narazaki K, Berg K, Stergiou N and Chen B. Physiological demands of competitive basketball. Scand J Med Sci Sport. 19: 425-322, 2009. https://www.ncbi.nlm.nih.gov/pubmed/18397196
  4. Deitch JR, Starkey C, Walters SL and Moseley JB. Injury Risk in Professional Basketball Players; A Comparison of Women’s National Basketball Association and National Basketball Association Athletes. Am J Sport Med. 34: 1077-1083, 2006. https://www.ncbi.nlm.nih.gov/pubmed/16493173
  5. Wertz, J., Galli, M., and Borchers, J. R. Achilles Tendon Rupture Risk Assessment for Aerial and Ground Athletes. Sport Health. 5: 407-409, 2013. https://www.ncbi.nlm.nih.gov/pubmed/24427410
  6. Caselli MA. Selecting the proper athletic shoe. Pod Manag. 25: 147-149, 2006.
  7. Commons AT and Low DC. Understanding the effect of high-cut shoes, running shoes and prophylactic supports on ankle stability when performing a v”-cut movement. Sport Exerc Med Open J. 1: 1-7, 2014.
  8. Sinclair, J. Effects of barefoot and barefoot inspired footwear on knee and ankle loading during running. Clin Biomech. 29: 395-399, 2014. https://www.ncbi.nlm.nih.gov/pubmed/24636307
  9. Sinclair, J., Atkins, S., Taylor, P. J., and Vincent, H. Effects of conventional and minimalist footwear on patellofemoral and Achilles tendon kinetics during netball specific movements. Comp Ex Phys. 11: 191-199, 2015.
  10. Sinclair, J., Hobbs, S. J., and Selfe, J. (2015). The Influence of Minimalist Footwear on Knee and Ankle Load during Depth Jumping. Research in Sports Medicine, 23(3), 289-301. https://www.ncbi.nlm.nih.gov/pubmed/26053415
  11. Sinclair, J., Edmundson, C.J., Brooks, D., and Hobbs, S.J. Evaluation of kinematic methods of identifying gait Events during running. Int J Sport Sci Eng. 5: 188-192, 2011.
  12. Thelen, D.G., Anderson, F.C., and Delp, S.L. Generating dynamic simulations of movement using computed muscle control. J Biomech. 36: 321–328, 2003. https://www.ncbi.nlm.nih.gov/pubmed/12594980
  13. Almonroeder, T., Willson, J.D., and Kernozek, T.W. The effect of foot strike pattern on Achilles tendon load during running. Annals Biomedical Eng. 41: 1758-1766, 2013. https://www.ncbi.nlm.nih.gov/pubmed/23640524
  14. Selvanetti, A.C.M., and Puddu, G. Overuse tendon injuries: basic science and classification. Op Tech Sport Med. 5: 110–17, 1997.
  15. Kirkendall, D.T., and Garrett W.E. Function and biomechanics of tendons. Scandinavian. J Med Sci Sport. 7: 62–66, 1997. https://www.ncbi.nlm.nih.gov/pubmed/9211605

Foot Posture Biomechanics and MASS Theory

by Edward S Glaser1 and David Fleming2*pdflrg

The Foot and Ankle Online Journal 9 (1): 12

Differences between single axis and postural models of foot biomechanics are explored.  Subtalar joint function is discussed and a new model explaining midfoot locking is proposed.  The posture of the foot is divided into zones of postural collapse.  The new postural model suggests that the foot’s posture controls its function and we should begin controlling the foot’s posture before postural collapse occurs.  Maximal Arch Supination Stabilization (MASS) posture is proposed as the geometry of a composite leaf spring that applies a calibrated, more evenly distributed, force per unit area opposing the postural collapse that occurs as the foot is intermittently compressed during ambulation.  One calibration method is explained.

Key words: biomechanics, posture, MASS theory

ISSN 1941-6806
doi: 10.3827/faoj.2016.0901.0012

1 – Founder and CEO of Sole Supports, Inc.
2 – Sole Supports, Inc.
* – Correspondence: dfleming@solesupports.com


Single Axis vs. Postural Biomechanics

The popular school of thought in foot biomechanics is a single axis approach. Merton Root was attempting to find something he could measure that would correlate to and predict deformity [1]. He discovered that by placing the patient prone while holding the off weight bearing foot in a palpated, “neutral” position, it was observed that most heels were inverted; rearfoot varus. He noticed both rearfoot and forefoot varus did correlate well to observations of deformities, lesions, and many lower extremity pathologies. Root recommended taking 17 measurements called the Static Biomechanical Exam [2]. Treatment was aimed at correcting what was viewed as a frontal plane deformity with frontal plane correction of the rearfoot and forefoot, called posts, designed to encourage the foot into a more neutral rotational position around the subtalar joint (STJ) axis. Excellent success was attained in most cases with the reduction of symptoms that gained this model of foot biomechanics broad acceptance.

At this writing, this model is still the backbone of the biomechanics curriculum of all colleges of podiatric medicine in the United States.  Merton Root, John Weed, and Bill Orien [2] did a thorough analysis of the motions that occur in the foot, analyzed muscle firing patterns, and did a magnificent job analyzing many of the most common, and therefore important, biomechanical foot deformities.  Neutral position, which Root defined as “neither pronated nor supinated”, is simply a rotational position around a singular axis; the subtalar joint axis.  Pronation and supination are defined in both the open and closed chains as rotations around this singular axis.  The extreme of single axis theory is to imagine that the foot only has one axis and consider the foot as just two rigid bodies teetering around this singular axis.  This model concerns itself with the distribution of kinetic forces and their perpendicular distance to this one axis.  This describes the Subtalar Axis Location and Rotational Equilibrium (SALRE) theory of Kevin Kirby, DPM [3].

The small amount of STJ rotation is where Merton Root and Kevin Kirby concentrated their attention [4].  According to Root’s own measurements the total range of STJ rotation in ideal gait is only six degrees (+2 to -4).  Pierrynowski, showed that palpation accuracy for “neutral” position with experienced practitioners is +/- 3 degree [5].  Meaning the best clinicians can find any position within the ideal range of motion of rotation around the STJ axis and call it “neutral”.

Craig Payne took this one step farther and studied the frontal plane rearfoot to forefoot variability of off weight bearing STJ neutral casting and found that new students, experienced doctors, and the peer-selected best caster all had the same frontal plane variation forefoot to rearfoot of 10 to 12 degrees within each group and just over 16 degrees between groups [6].  This frontal plane twist is the major determinant of arch height and, therefore, foot posture.  In a personal meeting with Dr. Payne, he said that they used his foot for all tests [7].   Subsequent examination of his foot revealed a fairly rigid forefoot in the frontal plane.  Sixteen degrees may have been his entire range of motion, and they found it all.

If any singular axis is to be chosen to describe foot biomechanics, the subtalar axis may be a particularly poor choice.  The reasoning behind this, is in physics, when a force is applied onto one side of an axis it causes rotation in one direction, as it moves to the other side of the axis rotation occurs in the opposite direction.  When the force passes directly through the axis, no rotational movement occurs.  The ground reactive force enters the foot ideally on the plantar posterior lateral aspect of the heel.  The STJ axis exits the foot at the same point; the momentum down the leg similarly passes its force vector down the center of the dome of the talus thereby intersecting the STJ axis.  The STJ axis is placed in an orientation that passes through the major forces entering the foot at heel contact, other than the force of friction which is horizontal and causes the forward roll of the calcaneus.

No singular axis can even begin to describe the motion that occurs during ambulation or simply the elevation and collapse of the arches of the foot. The foot has 26 bones and 35 joints, all of which move in some way.  Some of these joints are involved in rotation, and other joints simply slide in one plane.

Royal Whitman realized that the foot weakened its structure as its posture collapsed; calling the condition “weak foot” [8].  He may have been the first to have actually made the observation that foot posture controls its function.

Root et al, called Royal Whitman’s observation the phenomena of midtarsal locking and unlocking and attributed it to Elftman’s theory, that the talonavicular and calcaneocuboid axis deviated as the foot supinated [9].  Thus, this decreased the range of motion and parallelism of the axes, results in increased range of motion.  The talonavicular joint is an ovoid ball and socket having an infinite number of axes.  Sarrafian calls it the Acetabulum Pedis (hip socket of the foot) [10].  Whatever the rotational axis of the calcaneocuboid, the talonavicular joint will always find a parallel axis.

I propose that the locking mechanism of the midfoot is multifaceted.  When the talar head is directly on top to the anterior facet, sagittal plane motion between the talus and calcaneus is blocked.  Thus, when the gastroc-soleus complex fires, rotation occurs at the ankle joint.

Additionally, the Wring Theory by PC Jones describes twisting or wringing the foot into a more closed packed position; nesting the midfoot bones into each other [11].  Such a twisting would put more force on the first metatarsal head at toe off, per Root [1].  In a supinated posture the anterior facet of the STJ levels allowing transverse plane rotation of the talar head which carries with it the medial column of the foot which rides over the lateral column further restricting midfoot dorsiflexion.

Measurements taken of the geometry of the three facets on over 200 calcanei in The Terry Collection at the Smithsonian Institute yielded one consistent observation: when the anterior facet of the STJ is level in the frontal plane, the calcaneus is inverted [12].  Inversion occurs ideally at heel strike. This occurs at or near the end range of motion in subtalar supination.   The talocalcaneal motion, which is a posterior and slightly lateral slide along the cone-shaped posterior facet, is accompanied by a small amount of rotation around the STJ axis. This places the head of the talus squarely on the anterior facet.  This was also noted by Root [1].

The basic difference between single axis models, such as the STJ Neutral Model, and a postural model is that single axis models, by definition, ignore the rest of the foot.  You can find STJ neutral in a broad range of foot postures both in the open and closed kinetic chain.  Posture is simply stepping back and looking at the foot as a whole and observing the way elevation of the longitudinal arches causes bones to nest into each other in a more closed pack position.  Paul Jones attributes this to a generalized spiral twisting of the forefoot on the rearfoot, The Wring Theory [11].  Sarrafian described the frontal plane forefoot to rearfoot relationship as a twisted plate. All of these models are posture based [13].  Posture is the All Axis Model.

The foot is a machine with a tented structure.  The foot experiences intermittent compression between the downward force of the body and the ground, which is often in our society, a rigid surface like concrete or steel.  Clinical observation confirms that over a lifetime most individuals are genetically predisposed to postural collapse. Postural collapse loosens the foot’s structure and postural elevation tightens the foot. As Root [1] proposed, loosening allows for shock absorption and adaptation to the terrain and tightening prepares the foot for propulsion by creating a more rigid lever.

Postural Zones

What the STJ axis lacks in rotation, it more than makes up in translation.  If we observe the foot at heel strike through midstance, we see a huge forward and plantar grade translation of the STJ axis.  Southerland begins the Seven Theorems of Compensation in the Distal Human Lower Extremity with the words, “The foot hits the ground in a forward rolling motion” [14].  Jacqueline Perry describes the axis of translation as the heel rocker mechanism [15].  One of the more brilliant aspects of foot design is the round heel.  Like a ball, it has an infinite number of axes all passing through the center.   This allows us to hit the ground from any angle, forward or backward and apply the appropriate axis based on the direction of heel rotation with every step. Likewise, the STJ axis can translate through all of the following postural zones with each step.

Pathological Zone

Tom McPoil’s Tissue Stress Theory states that when microtrauma occurs faster than a person’s ability to heal, they experience a symptom [16].  During the last few degrees of postural collapse tissue stresses are highest.  Microtrauma occurring in this zone of foot posture causes symptoms.

Dysfunctional Zone

As the foot goes into further elevation of its posture, there is a zone where, according to Hammel, there is no significant rotation around the STJ axis in any plane [17]. Foot orthoses that attempt to elevate posture into this zone often cause medial longitudinal arch pain as the foot repeatedly drops down to impact the orthotic.  Hammel showed that from 25% to 90% of the stance phase of gait, no rotation in any plane occurs between the talus and the calcaneus.   The forefoot hits the floor at 27% of the way through the stance phase.  Ground reaction force applied to the forefoot displaces it superiorly in relation to the rearfoot. The most significant postural collapse occurs at this time.  Subtalar rotation in the transverse and sagittal planes occurs only from heel strike to 24% of stance.   Therefore, subtalar rotation and postural collapse are independent events occurring at different times in the gait cycle.  Early and excessive STJ rotation does, however, move the head of the talus off of the anterior facet loosening the foot’s structure, and preparing the foot for postural collapse.  Subtalar pronation is not synonymous with postural collapse, but it is a predicating factor.  Subtalar supination is not synonymous with postural elevation but is highly beneficial for efficient propulsion.  Pierrynowski and Trotter showed that elevation of the foot’s posture made a significant improvement in the Biomechanical Efficiency Quotient [18].

Functional Zone

As foot posture elevates beyond the Dysfunctional Zone the anterior facet of the STJ approaches level in the transverse plane.  This allows subtalar rotation to occur.  This is where the talar head slides posterior and rotates its six degrees around the STJ axis.  The closer the anterior facet is to level, the easier the subtalar rotation occurs and the rearfoot locks in the sagittal plane facilitating efficient propulsion.

Supination Instability Zone

Beyond the Functional Zone, there is a zone that is not always present, where the foot can be put into so much supination that it becomes laterally unstable.  As the downward force of the human body moves lateral to the foot, the propensity of inversion ankle sprain will increase due to a rotational moment created in that direction.

Composite Leaf Spring

Since the downward and deforming force of intermittent compression causes postural collapse, a corrective force would have to be applied in the opposite direction if functional change is desired.  This is where a foot orthotic device comes in.  A foot orthotic is a very simple machine.  It is a composite leaf spring.  There are two ways that such a leaf spring can be applied to the human foot.

Traditional orthotics based on the single axis models tend to be rather low in posture.  The cast is taken in a partially pronated position and then the arch is further lowered to varying degrees to make the orthotic tolerable.  Filling in, or lowering the arch of the orthotic, is often called “cast correction” even though it divorces the geometry of the foot from the geometry of the orthoses and allows for greater postural collapse before the orthotic is contacted by the arch.  Dysfunctional Zone postures are lowered to pathologic zone postures by arch fill.  As the foot reaches the end of its postural range of motion, ligaments are tightening up and the velocity of final impact is slowing down.  At this point the orthotic contacts the foot in the arch and the soft tissue compression dampen the final impact.  The Tissue Stress Model explains that symptoms are caused when microtrauma occurs faster than a person’s ability to heal [16].  Repetitive over stressing of the soft tissues can lead to bony malalignments that we refer to as foot deformities; hallux abducto valgus or hammertoes.  Symptoms become less evident when the amplitude of each tissue stressing event is decreased by soft tissue compression.  This dampening can mask symptoms without making a significant functional change in the gait cycle.  This explains an important contradiction.  These low, flat, smooth, invented, generic-shaped orthoses with their various tilts, skives, grooves, lumps, and bumps are simply herding the terminal tissue stresses around the bottom of the foot to mask symptoms with no appreciable change in kinematics.  Kirby et al, refers to these infinitesimal changes in kinematics that are so small as to be clinically meaningless [19].  He reports a statistically significant change in the angle of gait of less than 1.5 degrees, which is visually imperceptible.

This strategy is completely incapable of addressing posture because the foot is near its relaxed calcaneal stance posture when the foot’s medial longitudinal arch hits the orthotic. A corrective force applied after the motion has occurred can only mitigate the damage caused by the impact.  The orthotics are being used much like a car bumper.  Low velocity impacts cause little or no damage to the car because the bumper dampens the impact.

A different, and in this author’s opinion, better way to control the postural collapse of the foot would be to apply the corrective force throughout the entire gait cycle.  Simply choose a posture of the foot that approximates the beginning of the postural range of motion.  The spring flexes and limits the motion while continuously encouraging the foot back to its functional zone.  This is analogous to applying your brake and controlling the motion instead of mitigating the effect of repetitive impact.

MASS Posture

MASS posture has several elements.  First, it is the highest posture that the foot can attain at midstance, placing the foot in adequate supination to reach or approximate a level anterior facet of the STJ, putting it squarely within the functional zone.  The idea is simple.  If you want to control a motion, start at the beginning of that motion.

The foot poses a special problem.  The soft tissues between the orthotic and the bones compress unevenly. Therefore, an essential element of capturing the foot in this elevated posture is that the soft tissues must be evenly compressed as they will be during use.  There are many ways to achieve this.

A MASS Posture composite leaf spring applies an even distribution of force per unit of area by remaining in full contact with the foot throughout the gait cycle.  The foot never has to drop down to hit the orthotic because it is already touching it, which minimizes impact and thus tissue stresses.   It is the combination of full contact  (redistribution of force per unit area) eliminating hot spots and the lack of repetitive impact that allow such a spring to apply a rather large corrective force while remaining comfortable to most patients.  Once you have the correct geometry of the spring, it is time to adjust the spring constant.

Calibration

How much vertical force should this leaf spring apply to the foot in an evenly distributed manner?  Isaac Newton supplied the answer with his third law of motion:  for every action there is an equal and opposite reaction.  Applying that law to this problem:  the amount of force the orthotic should apply to the body is directly related to how much force the body is applying to the orthotic.  What causes the downward force of the human body onto the orthotic?

Obviously, body weight is a major factor.  Heavier people apply more force as measured by any household scale.  The more you weigh the greater the force the orthotic must resist and, therefore, the more rigid it must be.

Foot flexibility is another factor.  If the patient has Ehlers Danlos Disease, their ligaments are highly elastic and far less supportive.  They contribute little to the support of the foot’s posture and the orthotic must do more of the work.  If the patient’s ligaments are stiff, there is less range of motion and the foot generally collapses less.  The ligaments provide much of the support of the foot’s posture.  What little postural collapse occurs is easily elevated.  Foot flexibility can be measured in different ways.  One way to grade foot flexibility is to rotate the forefoot around the fifth metatarsal.  This is called the Gib Test or forefoot flexibility Forefoot Flexibility Test.  The foot can be graded from one to five [20].

Five, being the most rigid, is less than five degrees of total rotation of the forefoot on the rearfoot.  This can occur in Charcot foot, after a major trauma, or a surgical fusion.  A total rotation, up and down, between five and 30 degrees, tells us the foot is on the rigid side of normal and is graded a four.  Normal rotation is between 30 and 60 degrees and is graded a three.  The feet that are on the flexible side of normal will rotate between 60 and 85 degrees and are graded a two. The most flexible feet, that usually collapse the most, can rotate the forefoot around the fifth metatarsal more than 85 degrees and are graded a one.

This simple grading system is not meant to be accurate.  Accuracy is difficult to attain when the foot’s flexibility is a moving target and can change significantly throughout a single day.  Only an approximation is necessary or possible.  If the clinician is unsure of whether a patient is a one or three for example, it is best to report the smaller number.  That would tell the manufacturer that the foot is more flexible, and thus, make the orthotic more rigid.  If the device is actually too rigid for the patient, it can easily be recalibrated down to a more flexible spring by removing material, but as material is difficult to add accurately a spring made too flexible will have to be remade thicker from scratch.

Another way to assess forefoot flexibility more directly is to compare the actual curvature of the foot in its corrective MASS posture vs relaxed calcaneal versus stance posture.  These measurements are performed in the same medium to get comparable soft tissue compression in both casts.  Comparing the best and worst posture will help the patient and clinician understand how they would benefit from influencing the foot’s posture.   This is far more objective because it shows the actual change in curvature that will occur with the posture is collapsed versus restored.

Momentum (mass times velocity) is the third factor that affects the magnitude of the downward force of the body.  Running over a force plate produces more impact force than walking. Therefore, we must consider a range of forces to resist called, ADL or activities of daily living, and calibrate the orthotic to deliver an equal and opposite range.  Athletes may have a different range of forces, these can be referred to as training or competing ranges, which are much higher.  A power lifter, for example, may want an orthotic calibrated to resist his entire weight plus the weight he is deadlifting or squatting.  That same athlete will need a different pair of orthotics for his ADL.

Munteanu showed that the more elevated the foot’s posture is, the less supination resistance is measured, and the more collapsed the foot’s posture is, the greater the force necessary to elevate the arch [21].  Postural elevation makes further elevation easier and postural collapse makes further postural collapse easier.   Several lever arms in the foot increase and decrease to accomplish this.

Measuring the upward force delivered by the orthotic is difficult.  Calibration of the orthotic can be accomplished by the application of Pascal’s Law; pressure inside an enclosed container is equivalent in all directions.  Place the orthotic in an enclosed container and blow up a bladder over the orthotic.  As the bladder expands, it fully contacts the orthotic and begins to flex it.   Flexion can then be captured digitally either via liner encoder or optics.  A force curve plotting flexion against pressure gives us a slope.  This slope correlates to the spring constant, which allows each orthotic to be calibrated.

Root found that in ideal gait 60% of the force applied to the ground at toe off should be under the first ray [22]. Higby measured the force distribution on the metatarsal heads at toe off [23]. What are these forces? Initially, MASS posture orthotics transferred 44% more force to the first metatarsal head at toe off than neutral position orthotics with posts. At six weeks this difference grew to 61% (p=.006) [24].  This means that when the arch is raised, the first ray not only comes down and lateral, but additionally increases its purchase.

Conclusion

Posture controls function.  Postural collapse is the cause of functional impairment in the majority of foot patients and often leads to pain, pathology, and deformity.  MASS posture is an aggressive approach to foot biomechanics.  It attempts to restore as close to an ideal posture to the foot as each foot can tolerate with its individual anatomy.   Application of a calibrated leaf spring to resist collapse of foot posture can often make early visible changes in the gait cycle.  A composite leaf spring, or orthotic, must begin to resist pathologic motion or postural collapse before that motion occurs.  MASS posture theory is proposed, which is a plastic leaf spring that is in full contact with the foot in the highest posture that a person can attain at mid-stance with the heel and forefoot in contact with the ground and the soft tissues evenly compressed.  Such a spring calibrated to deliver an equal and opposite range of forces to those applied by the body, encourages the foot into a more functional foot posture that may reverse deformity.  Form follows pathological function in the direction of disease and deformity.   It stands to reason that it would similarly follow restored function in the direction of health and reversal of deformity. Further research is needed to determine the measurable effects on several diagnoses and to explore better ways to measure and document the gait changes achieved by MASS Posture.

References

  1. Root, M. L., Orien, W. P., & Weed, J. H. (1977).Normal and abnormal function of the foot (Vol. 2). Los Angeles: Clinical Biomechanics Corporation.
  2. Root, M. I. (1973). Biomechanical examination of the foot.Journal of the American Podiatry Association63(1), 28-29.
  3. Kirby, K. A. (2001). Subtalar joint axis location and rotational equilibrium theory of foot function.Journal of the American Podiatric Medical Association91(9), 465-487.
  4. Kirby, K. A. (2010). Evolution of foot orthoses in sports. InAthletic Footwear and Orthoses in Sports Medicine (pp. 19-35). Springer New York.
  5. Pierrynowski, M. R., Smith, S. B., & Mlynarczyk, J. H. (1996). Proficiency of foot care specialists to place the rearfoot at subtalar neutral.Journal of the American Podiatric Medical Association86(5), 217-223.
  6. Chuter, V., Payne, C., & Miller, K. (2003). Variability of neutral-position casting of the foot.Journal of the American Podiatric Medical Association93(1), 1-5.
  7. According to Dr. C. Payne (Personal Communication)
  8. Whitman, R. (2010). The Classic: A Study of the Weak Foot, with Reference to its Causes, its Diagnosis, and its Cure; with an Analysis of a Thousand Cases of So-Called Flat-Foot.Clinical Orthopaedics and Related Research®468(4), 925-939.
  9. Elftman, H. (1960). The transverse tarsal joint and its control.Clinical orthopaedics16, 41.
  10. Sarrafian, S. K. (1993). Biomechanics of the subtalar joint complex.Clinical orthopaedics and related research290, 17-26.
  11. Jones, P. C. (2011). Unwringing the Helix.Podiatry Management30(7).
  12. Glaser ES, Fleming DC, Reece N. Interval measurement of the angle of calcaneal facets: A historical postmortem study. Foot Ankle Online J. 2016, 9(1):8.
  13. Sarrafian, S. K. (1987). Functional characteristics of the foot and plantar aponeurosis under tibiotalar loading.Foot & Ankle International8(1), 4-18.
  14. Southerland CC, Orien WP (1995). Seven theorems of compensation in the distal human lower extremity. The Lower Extremity 2, (3), 1
  15. Perry, J., & Burnfield, J. M. (1993).Gait analysis: normal and pathological function. Slack.
  16. McPoil, T. G., & Hunt, G. C. (1995). Evaluation and management of foot and ankle disorders: present problems and future directions.Journal of Orthopaedic & Sports Physical Therapy21(6), 381-388.
  17. Hamel, A. J., Sharkey, N. A., Buczek, F. L., & Michelson, J. (2004). Relative motions of the tibia, talus, and calcaneus during the stance phase of gait: a cadaver study.Gait & posture20(2), 147-153.
  18. Trotter, L. C., & Pierrynowski, M. R. (2008). Changes in Gait Economy Between Full-Contact Custom-made Foot Orthoses and Prefabricated Inserts in Patients with Musculoskeletal Pain A Randomized Clinical Trial.Journal of the American Podiatric Medical Association98(6), 429-435.
  19. Huerta, J. P., Moreno, J. M. R., Kirby, K. A., Carmona, F. J. G., & García, A. M. O. (2009). Effect of 7-degree rearfoot varus and valgus wedging on rearfoot kinematics and kinetics during the stance phase of walking.Journal of the American Podiatric Medical Association99(5), 415-421.
  20. Glaser, E., Bursch, D., & Currie, S. J. (2006). Theory, practice combine for custom orthoses.Biomechanics13(9), 33-43.
  21. Munteanu, S., Bassed, A., & Payne, C. (2003). Supination resistance, Foot Posture Index and effect of foot orthoses on first MPJ range of motion.
  22. Root, M. L., Orien, W. P., & Weed, J. H. (1977). Functions of the muscles of the foot.Normal and Abnormal Function of the Foot, pp250–252, edited by ML Root, WP Orien, JH Weed, Clinical Biomechanics Corporation, Los Angeles.
  23. Hodgson, B., Tis, L., Cobb, S., McCarthy, S., & Higbie, E. (2006). The Effect of 2 Different Custom-Molded Corrective Orthotics on Plantar Pressure.Journal of Sport Rehabilitation15(1).
  24. Hodgson, B., Tis, L., Cobb, S., McCarthy, S., & Higbie, E. (2006). The effect of 2 different custom-molded corrective orthotics on plantar pressure.Journal of Sport Rehabilitation15(1), 33.

 

 

 

The influence of barefoot and shod running on Triceps surae muscle strain characteristics

by Sinclair J1*, Cole T2, Richards J2pdflrg

The Foot and Ankle Online Journal 9 (1): 4

The aim of the current investigation was to determine the effects of barefoot and shod running on the kinematics of the Triceps-Surae muscle group. Twelve male participants ran at 4.0 m.s-1 (± 5%) in both barefoot and shod conditions. Kinematics were measured using an eight-camera motion analysis system. Muscle kinematics from the lateral Gastrocnemius, medial Gastrocnemius and Soleus were obtained using musculoskeletal modelling software (Opensim v3.2).  The results showed that muscle strain for the lateral Gastrocnemius (barefoot = 1.10 & shod = 0.33 %), medial Gastrocnemius (barefoot = 1.07 & shod = 0.32 %) and Soleus (barefoot = 3.43 & shod = 2.18 %) were significantly larger for the barefoot condition. Given the proposed association between the extent of muscle strain and the etiology of chronic muscle strain pathologies, the current investigation shows that running barefoot may place runners at greater risk from Triceps-Surae strain injuries.

Key words: Biomechanics, barefoot, shod, Triceps-Surae

ISSN 1941-6806
doi: 10.3827/faoj.2016.0901.0004

1 – Centre for Applied Sport Exercise and Nutritional Sciences, School of Sport & Wellbeing, College of Health & Wellbeing, University of Central Lancashire, UK.
2 – Allied Health Research Unit, School of Health, College of Health & Wellbeing, University of Central Lancashire, UK.
* Correspondence: Dr. Jonathan Sinclair, jksinclair@uclan.ac.uk


Engaging in recreational and competitive distance running has been shown to provide a number of health benefits [1]. Despite this runners are highly susceptible to chronic injuries [2], with an occurrence rate of around 80 % over the course of one year [3]. A large number of strategies have been investigated in biomechanical research with the specific aim of attenuating the risk of running injuries.

One such conservative strategy is to choose running shoes with appropriate mechanical characteristics; the properties of running shoes have been proposed as a mechanism by which chronic injuries can be controlled [4]. Recently barefoot running has been the focus of much attention in biomechanics research.

The popularity and attention paid to barefoot footwear is due the proposition that running barefoot may be able to reduce the incidence of chronic running injuries [5, 6].

The findings from biomechanical research into the kinetics and kinematics of running barefoot in comparison to shod have been equivocal. Sinclair et al. [7] examined the effects of barefoot and shod running on kinetics, kinematics and tibial accelerations during the stance phase. Their kinematic observations showed that the ankle was significantly more plantarflexed at footstrike in the barefoot condition. In addition it was also shown the running barefoot was associated with significantly greater tibial accelerations and vertical rates of loading. Sinclair et al. [8] similarly investigated the effects of barefoot and shod conditions on running kinetics and kinematics. Their kinematic findings showed that barefoot running was associated with a more plantarflexed ankle position at footstrike and also a greater peak eversion angle. The kinetic findings indicated that barefoot running demonstrated a significantly greater vertical rate of loading. When comparing the kinetics and sagittal plane kinematics of running barefoot and shod, Lieberman et al [5] demonstrated firstly that the ankle was significantly more plantarflexed at footstrike in the barefoot condition. However, their kinetic observations showed that the vertical rate of loading was larger when running with shoes. Similarly, Squadrone & Gallozzi, [9] showed that running barefoot was associated with increased plantarflexion at footstrike but with subsequent reductions in peak vertical impact forces.

In addition, with the development more accurate musculoskeletal models more recent research has been able to investigate the loads experienced by specific musculoskeletal structures. Bonacci et al, [10] showed that running barefoot was associated with significant reductions in patellofemoral loading in comparison to shod. Sinclair, [11] similarly demonstrated that patellofemoral loading was significantly reduced when running barefoot but that running without shoes mediated subsequent increases in the loads borne by the Achilles tendon. Finally, Sinclair et al, [12] investigated the effects of barefoot and shod running on limb and joint stiffness characteristics during the stance phase. They showed that limb and knee stiffness were greater when running barefoot but that ankle stiffness was greater when running shod.

There is currently a paucity of biomechanical research investigating muscle mechanics during barefoot and shod running. Sinclair et al, [13] investigated the effects of barefoot and shod running on lower limb muscle forces during the stance phase of running. Their observations showed that peak forces from the Rectus femoris, Vastus medialis, Vastus lateralis and Tibialis anterior were significantly larger in the shod condition whereas Gastrocnemius forces were significantly larger during barefoot running. Similarly, Sinclair, [14] studied the effects of running barefoot and shod on peak and mean foot muscle forces. The findings confirmed that peak and mean forces from the Flexor digitorum longus, Flexor hallucis longus, Peroneus longus muscles were significantly larger when running barefoot, whereas peak and average forces of the Extensor digitorum longus and Extensor hallucis longus muscles were significantly larger when running shod.

There has yet to be any published research investigating Triceps Surae muscle mechanics during barefoot and shod running. Anecdotal evidence of calf pain and stiffness has been reported by runners who seek to conduct their training without shoes. Furthermore, the prospective investigation of Altman & Davis [15] showed that calf injuries may be more prominent in barefoot runners in comparison to those who train shod. This indicates that an investigation into the mechanics of the Tricep-surae (calf) muscle group during barefoot and shod running would be of both practical and clinical significance to both clinicians and runners themselves.

Therefore the aim of the current investigation was to determine the effects of barefoot and shod running on the kinematics of the Triceps Surae muscle group. A study of this nature may aid our understanding of muscle function during barefoot running. The current work tests the hypothesis that the magnitude of strain experienced by the Triceps Surae muscles will be significantly larger when running barefoot.

Methods

Participants

Twelve male runners (age 23.58 ± 2.88 years, height 1.77 ± 0.10 cm and body mass 79.40 ± 5.87 kg) volunteered to take part in this study. All runners were free from musculoskeletal pathology at the time of data collection. Participants provided written informed consent in accordance with the principles outlined in the Declaration of Helsinki. Each runner was considered to be exhibit a natural rearfoot strike pattern as they exhibited an impact peak in their vertical ground reaction force curve when wearing conventional footwear. The procedure was approved by the University of Central Lancashire ethical committee.

Procedure

Participants ran at a velocity of 4.0 m.s-1 ±5%, striking an embedded force platform (Kistler, Kistler Instruments Ltd., Alton, Hampshire) with their right (dominant) foot [16]. The velocity of running was monitored using infrared timing gates (Newtest, Oy Koulukatu, Finland). The stance phase was defined as the duration over which 20 N or greater of vertical force was applied to the force platform [17]. All runners completed five successful trials in each footwear condition.

Kinematic information was captured at 250 Hz using an eight camera optoelectric motion analysis system (QualisysTM Medical AB, Goteburg, Sweden). To define the anatomical frames of the trunk, pelvis, thighs, shanks and feet retroreflective markers were placed at the C7, T12 and xiphoid process landmarks and also positioned bilaterally onto the acromion process, iliac crest, anterior superior iliac spine, posterior super iliac spine, medial and lateral malleoli, medial and lateral femoral epicondyles and greater trochanter. Carbon-fiber tracking clusters comprising of four non-linear retroreflective markers were positioned bilaterally onto the thigh and shank segments. 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.

Data processing

Marker trajectories were filtered 12 Hz using a low pass Butterworth 4th order zero-lag filter and analyzed using Visual 3D (C-Motion, Germantown, MD, USA. All information was normalized to 100 % of the stance phase. For the current study angular kinematics of the ankle joint were examined. Kinematic measures from the ankle were extracted for statistical analysis were 1) angle at footstrike and 2) relative peak range of motion from footstrike to peak angle.

OpenSim software was used to quantify muscle-tendon lengths during the stance phase of running [18]. Muscle kinematics were quantified using the gait2392 model using OpenSim v3.2. This model corresponds to the eight segments exported from Visual 3D and features ninety two muscles, eighty six of which are centered around the lower extremities and six are associated with the pelvis and trunk. The muscle properties were modelled using the Hill recommendations based on the associations between force-velocity-length [19]. These muscle properties were then scaled based on each participant’s height and body mass based on the recommendations of Delp et al, [20]. Muscle-tendon lengths are determined by the positions of their proximal and distal muscles muscle origins. The muscle–tendon units which were evaluated as part of the current research were the lateral Gastrocnemius, medial Gastrocnemius, and Soleus. Muscle kinematic parameters that were extracted for statistical analysis were 1) eccentric strain (representative of the maximum increase in muscle length divided by the length at footstrike and 2) peak lengthening velocity.

In addition to this we also estimated the total muscle strain experienced per mile (% x mile) by multiplying the muscle strain magnitude by the number of steps required to complete one mile. The number of steps required to complete one mile was calculated using the step length. Step length was obtained by taking the difference in the horizontal position of the foot between the right and left legs at footstrike [21, 22].

Statistical analyses

Descriptive statistics (means, standard deviations and 95% confidence intervals) were obtained for each footwear condition. Shapiro-Wilk tests were used to screen the data for normality. Footwear mediated differences in foot muscle kinetics were examined using paired samples t-tests. All statistical actions were conducted using SPSS v22.0 (SPSS Inc, Chicago, USA).

Results

Figures 1-3 and table 1 show ankle joint and muscle kinematics as a function of barefoot and shod running conditions. The results show that the different running conditions significantly influence both joint and muscle kinematics.

Ankle kinematics

The ankle was found to be significantly (t (11) = 4.51, p<0.05) more plantarflexed at footstrike in the barefoot conditions in comparison to shod. Furthermore, the relative range of motion was found to be significantly (t (11) = 4.08, p<0.05) greater when running barefoot in comparison to shod (Figure 1).

Muscle kinematics

For the lateral Gastrocnemius muscle running barefoot was associated with significantly (t (11) = 2.81, p<0.05) larger muscles strain in comparison to shod running (Figure 2a; Table 1). In addition when running barefoot the lateral Gastrocnemius exhibited a significantly (t (11) = 2.37, p<0.05) greater lengthening velocity than during shod running (Figure 2a; Table 1). Finally barefoot running was associated with a significantly (t (11) = 2.81, p<0.05) greater strain experienced per mile (Table 1).

For the medial Gastrocnemius muscle running barefoot was associated with significantly (t (11) = 2.79, p<0.05) larger muscle strain in comparison to shod running (Figure 2b; Table 1). In addition when running barefoot the medial Gastrocnemius exhibited a significantly (t (11) = 2.39, p<0.05) greater lengthening velocity than during shod running (Figure 3b; Table 1). Finally barefoot running was associated with a significantly (t (11) = 2.83, p<0.05) greater strain experienced per mile (Table 1).

For the Soleus muscle running barefoot was associated with significantly (t (11) = 3.79, p<0.05) larger muscle strain in comparison to shod running (Figure 2c; Table 1). In addition when running barefoot the Soleus exhibited a significantly (t (11) = 2.69, p<0.05) greater lengthening velocity than during shod running (Figure 3c; Table 1). Finally barefoot running was associated with a significantly (t (11) = 3.93, p<0.05) greater strain experienced per mile (Table 1).

Fig1

Figure 1 Sagittal ankle kinematics as a function of barefoot and shod conditions (black = barefoot and grey = shod) (DF = dorsiflexion).

Fig2

Figure 2 Tirceps Surae muscle kinematics as a function of barefoot and shod conditions (black = barefoot and grey = shod) (a. = lateral Gastrocnemius, b. = medial Gastrocnemius, c. = Soleus).

Fig3

Figure 3 Tirceps Surae muscle velocities as a function of barefoot and shod conditions (black = barefoot and grey = shod) (a. = lateral Gastrocnemius, b. = medial Gastrocnemius, c. = Soleus).

Tab1

Table 1 Triceps Surae muscle kinematics (Means, SD’s & 95% CI’s) as a function of barefoot and shod conditions.

Discussion

The aim of the current investigation was to quantify the effects of barefoot and shod running on Triceps Surae muscle kinematics. To the authors knowledge this represents the first comparative analysis of Triceps Surae mechanics when running in different footwear.

The first key observation from the current paper is that ankle was shown to be significantly plantarflexed at footstrike in the barefoot condition in comparison to running shod. This indicates that runners modified their footstrike pattern and adopted a non-rearfoot strike when running barefoot. This finding concurs with the observations of Squadrone & Gallozzi, [9], Lieberman et al, [5] and Sinclair et al, [7, 8] who each showed a more plantarflexed ankle position when wearing running barefoot. It proposed that this finding relates to the absence of shoe cushioning when running barefoot. Runners adopt a non-rearfoot strike pattern in order to compensate for the lack of a shoe midsole and attenuate the loads experienced by the musculoskeletal system [5]. The first key finding from the current work is that strain magnitude and velocity in each of the three muscles associated with the Triceps-Surae was significantly larger in the barefoot condition in comparison to shod. This observation supports our original hypothesis and may have clinical significance. Muscle strains occur as a function of excessive muscle lengthening during periods of eccentric muscle lengthening [23]. The findings from the current investigation therefore support the proposition of Altman & Davis, [15] in that running barefoot appears to place runners at increased risk from Triceps-Surae strain injuries.

It is proposed that these observations relate to the change in footstrike pattern and increased range of motion mediated by running without shoes. The Triceps-Surae muscles insert distally into the Achilles tendon insertion and proximally at the posterior aspects of the tibia/ femur. Therefore the increased plantar flexion at footstrike observed when running barefoot means that the muscles are in a shortened position compared shod running. This in conjunction with the increased dorsiflexion range of motion at the ankle means that the Triceps-Surae must lengthen to a greater extent given the anterior translation of the proximal muscle insertion points. This finding therefore suggests that whilst the non-rearfoot strike pattern associated with barefoot running may reduce the load experienced by the patellofemoral joint [10, 11] and also vertical rate of loading [5,9] it may be at the expense of increased Triceps-Surae strain.

The findings in relation to muscle strains from the current investigation can be further contextualized taking into account the increased number of steps required to complete one mile when running barefoot. This led to further increases in the amount of muscle strain experienced per mile, over and above those reported per footfall when participants ran barefoot. Therefore, whilst the amount of strain experienced per footfall is relatively small when contrasted against muscle strains shown in other sports [24], because running represents a cyclical activity which involves multiple footfalls the cumulative strain is high. This observation further supports the notion that running barefoot may enhance the likelihood of experiencing a chronic muscle strain injury at the Triceps-Surae.

In conclusion, although differences in the effects of barefoot running have been examined extensively, the current knowledge regarding the differences in Triceps-Surae kinematics between barefoot and shod running is limited. The present investigation therefore adds to the current knowledge by providing a comprehensive evaluation of Triceps-Surae muscle kinematic parameters when running in barefoot and shod conditions. On the basis muscle strain parameters were significantly greater when running barefoot; the findings from the current investigation indicate that barefoot running may place runners at increases risk from chronic Triceps-Surae muscle strain injuries in comparison to running shod.

References

  1. Schnohr P, O’Keefe JH, Marott JL, Lange P, Jensen GB. Dose of jogging and long-term mortality: the Copenhagen City Heart Study. Journal of the American College of Cardiology 2015; 65: 411-419. (PubMed)
  2. Taunton JE, Clement DB, McNicol K. Plantar fasciitis in runners. Canadian Journal of Applied Sport Sciences 1982; 7: 41-44. (PubMed)
  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. (PubMed)
  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. Lieberman DE, Venkadesan M, Werbel WA, Daoud AI, D’Andrea S, Davis IS, et al. Foot strike patterns and collision forces in habitually barefoot versus shod runners. Nature; 2010; 463: 531-535. (Link)
  6. Warburton, M. Barefoot running. Sportscience 2000; 5; 1-4.
  7. Sinclair J, Greenhalgh A, Brooks D, Edmundson CJ, Hobbs SJ. The influence of barefoot and barefoot-inspired footwear on the kinetics and kinematics of running in comparison to conventional running shoes. Footwear Science 2013: 5, 45-53.
  8. Sinclair, J, Hobbs, SJ, Currigan, G, Taylor PJ. A comparison of several barefoot inspired footwear models in relation to barefoot and conventional running footwear. Comparative Exercise Physiology 2013; 9: 13-21.
  9. Squadrone R, Gallozzi C. Biomechanical and physiological comparison of barefoot and two shod conditions in experienced barefoot runners. Journal of Sports Medicine & Physical Fitness 2009; 49: 6-13. (PubMed)
  10. Bonacci J, Vicenzino B, Spratford W, Collins P. Take your shoes off to reduce patellofemoral joint stress during running. British Journal of Sports Medicine, (In press). (Link)
  11. Sinclair J. Effects of barefoot and barefoot inspired footwear on knee and ankle loading during running. Clinical Biomechanics 2014; 29: 395-399. (PubMed)
  12. Sinclair J, Atkins, S, Taylor PJ. The Effects of Barefoot and Shod Running on Limb and Joint Stiffness Characteristics in Recreational Runners. Journal of Motor Behavior 2015 (In press). (PubMed)
  13. Sinclair J, Atkins S, Richards J, Vincent H. Modelling of Muscle Force Distributions During Barefoot and Shod Running. Journal of Human Kinetics 2015 (In press).
  14. Sinclair, J. (2015). Barefoot and shod running: their effects on foot muscle kinetics. FAOJ 2015; 8: 2. (Link)
  15. Altman AR, Davis IS. Prospective comparison of running injuries between shod and barefoot runners. British Journal of Sports Medicine, 2015 (In press). (Link)
  16. 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 30: 166–172. (PubMed)
  17. Sinclair J, Edmundson CJ, Brooks D, Hobbs SJ. Evaluation of kinematic methods of identifying gait Events during running. International Journal of Sport Science & Engineering 2011; 5: 188-192. (Link)
  18. Delp SL, Anderson FC, Arnold AS, Loan P, Habib A, John CT, Thelen DG. OpenSim: open-source software to create and analyze dynamic simulations of movement. IEEE Transactions on Biomedical Engineering 2007; 54: 1940-1950. (PubMed)
  19. Zajac FE. Muscle and tendon: properties, models, scaling, and application to biomechanics and motor control. Critical Reviews in Biomedical Engineering. 1989; 17: 359–411.
  20. Delp SL, Loan JP, Hoy MG, Zajac FE, Topp EL, Rosen JM. An interactive graphics-based model of the lower extremity to study orthopaedic surgical procedures. IEEE Transactions on Biomedical Engineering 1990; 37: 757–767. (PubMed)
  21. Almonroeder, T, Willson, JD, Kernozek, TW. The effect of foot strike pattern on Achilles tendon load during running. Annals of Biomedical Engineering 2013; 41: 1758-1766. (PubMed)
  22. Sinclair J, Richards J, Shore H. Effects of minimalist and maximalist footwear on Achilles tendon load in recreational runners. Comparative Exercise Physiology 2015 (In press).
  23. Mueller-Wohlfahrt HW, Haensel L, Mithoefer K, Ekstrand J, English B, McNally S, Ueblacker P. Terminology and classification of muscle injuries in sport: a consensus statement. British Journal of Sports Medicine 2012; 47: 342-350. (PubMed)
  24. Sinclair J. Side to side differences in hamstring kinematics during maximal instep kicking in male soccer players. Movement & Sport Sciences 2015 (In press).

Effects of foot orthoses on patellofemoral load in recreational runners

by Sinclair J1, Vincent H1, Selfe J2, Atkins S1, Taylor PJ3, and Richards J2pdflrg

The Foot and Ankle Online Journal 8 (2): 5

The most common chronic injury in recreational runners is patellofemoral pain. Whilst there is evidence to suggest that orthotic intervention may reduce symptoms in runners who experience patellofemoral pain the mechanism by which their clinical effects are mediated is currently poorly understood. The aim of the current investigation was to determine whether foot orthoses reduce the loads experienced by the patellofemoral joint during running. Patellofemoral loads were obtained from fifteen male runners who ran at 4.0 m·s-1. Patellofemoral loads with and without orthotics were contrasted using paired t-tests. The results showed that patellofemoral joint loads were significantly reduced as a function of running with the orthotic device. The current investigation indicates that through reductions in patellofemoral loads, foot orthoses may serve to reduce the incidence of chronic running injuries at this joint.

Keywords: patellofemoral pain, orthoses, biomechanics

ISSN 1941-6806
doi: 10.3827/faoj.2015.0802.0005

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

1. Division of Sport Exercise and Nutritional Sciences, School of Sport Tourism and Outdoors, University of Central Lancashire.
2. Allied Health Research Unit, School of Sport Tourism and Outdoors, University of Central Lancashire.
3. School of Psychology, University of Central Lancashire.


D
istance running has been shown to be physiologically beneficial [1]. However despite this, research examining the incidence of running injuries indicates that chronic pathologies are a prominent complaint for both recreational and competitive runners [2], with an incidence rate of around 70% during the course of a year [3].

The most common chronic injury in recreational runners is patellofemoral pain, which is characterized by pain linked to the contact of the posterior surface of the patella with the femur during dynamic activities [4].

Pain symptoms, which develop as a function of patellofemoral disorders can be debilitating and patellofemoral pain may also be a pre-cursor to the progression of osteoarthritis in later life [5,6]. Conservative treatment of patellofemoral disorders is preferable to operative interventions, and the efficacy of a number of conservative approaches has been explored in the literature.

There is evidence to suggest that orthotic intervention may reduce symptoms in runners who experience patellofemoral pain. Collins et al. prospectively examined the efficacy of foot orthoses in the management of patellofemoral pain [7]. Foot orthoses were shown to produce clinically meaningful improvements in pain symptoms. Eng et al. examined the effectiveness of soft foot orthotics in the treatment of patients with patellofemoral pain syndrome [8]. Participants were assigned to either an orthotic or control condition and subjects reported their perceived pain levels over an 8-week period using a visual analogue scale. It was shown that the soft foot orthotics may be an effective treatment mechanism for patellofemoral pain. Batron et al. investigated the effects of 12-week intervention using of non-custom foot orthoses on self-reported improvements in pain symptoms [9]. It was shown that 25% of participants showed marked improvements in patellofemoral pain symptoms as a function of orthotic intervention. Pitman & Jack monitored the efficacy of foot orthoses as a treatment modality for patellofemoral pain [10]. They found that orthotics produced reductions in pain symptoms, which led to the conclusion that orthotics may be an effective treatment mechanism.

Despite the potential efficacy of foot orthoses in the prevention/treatment of patellofemoral pain symptoms, there is a paucity of research investigating any potential alterations in loading at this joint that may be mediated through orthotic intervention. The aim of the current investigation was therefore to determine whether foot orthoses reduce the loads experienced by the patellofemoral joint during the stance phase of running. This study tests the hypothesis that orthoses will reduce patellofemoral load during running.

Methods

Participants

Fifteen male participants (Age 25.76 ± 5.21 years; height 1.74 ± 0.06 m; mass 71.15 ± 4.84 kg) took part in the current study. Participants were all recreational runners who engaged in training at least three times per week. Ethical approval for this project was obtained from the University and each participant provided informed consent in written form in accordance with the declaration of Helsinki.

Orthotic device

Commercially available orthotics (Sorbothane, shock stopper sorbo Pro; Nottinghamshire UK) were examined in the current investigation. Although the right side was selected for analysis orthotic devices were placed inside both shoes.

Procedure

Participants completed five trials running at 4.0 m·s-1 with and without orthotics. The order in which participants ran in each condition was counterbalanced. Participants ran over an embedded piezoelectric force platform (Kistler Instruments, Model 9281CA) operating at 1000 Hz [11]. Running velocity was controlled using infrared timing gates (SmartSpeed Ltd UK). A deviation of ±5% from the pre-determined velocity was allowed. Participants struck the force platform with their right (dominant) limb and five trials were obtained from each footwear condition. Three-dimensional (3-D) kinematics and ground reaction forces data were collected synchronously. The stance phase was defined as the duration over which >20 N of vertical force was applied to the force platform [12]. Kinematic information was obtained using an eight camera optoelectric motion capture system (Qualisys Medical AB, Goteburg, Sweden) using a capture frequency of 250 Hz. Dynamic calibration of the motion capture system was conducted prior to data collection.

The current investigation used the calibrated anatomical systems technique (CAST) to model the lower extremity segments in six degrees of freedom [13]. To define the anatomical frame of the right shank and thigh, retroreflective markers were positioned unilaterally to the medial and lateral malleoli, medial and lateral epicondyle of the femur and greater trochanter. Rigid technical tracking clusters were positioned on the shank and thigh segments. Static trials were conducted 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.

Data processing

Ground reaction force and kinematic data were smoothed using cut-off frequencies of 50 Hz and 12 Hz with a low-pass Butterworth 4th order filter using Visual 3-D (C-Motion, Germantown, MD, USA). Newton-Euler inverse-dynamics were used which allowed knee joint moments to be calculated. Knee loading was examined through extraction of peak knee extensor moment, peak knee abduction moment, patellofemoral contact force (PTCF) and patellofemoral contact pressure (PTCP).

A previously utilized algorithm was used to quantify PTCF and PTCP [14]. This method has been utilized previously to resolve differences in PTCF and PTCP when using different footwear [15,16,17] and between those with and without patellofemoral pain [18]. PTCF (B.W) was estimated using knee flexion angle (KFA) and knee extensor moment (KEM) through the biomechanical model of Ho et al [19]. The moment arm of the quadriceps (QMA) was calculated as a function of KFA using a non-linear equation, based on cadaveric information presented by van Eijden et al. [20]:

QMA = 0.00008 KFA 3 – 0.013 KFA 2 + 0.28 KFA + 0.046

Quadriceps force (FQ) was calculated using the below formula:

FQ = KEM / QMA

PTCF was estimated using the FQ and a constant (C):

PTCF = FQ C

The C was described in relation to KFA using the equation described by van Eijden et al. [20]:

C = (0.462 + 0.00147 KFA 2 – 0.0000384 KFA 2) / (1 – 0.0162 KFA + 0.000155 KFA 2 – 0.000000698 KFA 3)

PTCP (MPa) was calculated using the PTCF divided by the patellofemoral contact area. The contact area was delineated by fitting a 2nd-order polynomial curve to the data of Powers et al., [21] showing patellofemoral contact areas at varying levels of KFA.

PTCP = PTCF / contact area

PTCF loading rate (B.W/s) was also calculated as a function of the change in PTFC from initial contact to peak force divided by the time to peak force.

Statistical Analyses

The data were tested for normality using a Shapiro-Wilk test which confirmed that the data were suitable for parametric testing. Means and standard deviations were calculated for each running condition. Differences in the outcome 3D kinematic parameters were examined using paired samples t-tests. The alpha level required for statistical significance was adjusted to p≤0.008 based on the number of comparisons being made. Effect sizes for all significant observations were calculated using a Cohen’s D statistic. All statistical analyses were conducted using SPSS v21.0 (SPSS Inc, Chicago, USA).

Results

 fig1

Figure 1 Knee kinetics and kinematics as a function of orthotic intervention, black = no-orthotic and dash = orthotic, (a= knee angle, b = sagittal knee moment c = PTCF, d = PTCP, e = coronal knee moment) (FL = flexion, EX = extension, AD = adduction).

Peak knee extensor moment was significantly (t (14) = 4.11, p<0.008, D = 2.20) greater in the non-orthotic condition compared to running with orthotics (Table 1, Figure 1a). In addition PTFC (t (14) = 3.96, p<0.008, D = 2.12) and PTCP (t (14) = 4.57, p<0.008, D = 2.44) were also shown to be significantly greater in the non-orthotic condition compared to running with orthotics (Table 1, Figure 1bc). Finally PTCF loading rate was shown to be significantly (t (14) = 3.88, p<0.008, D = 2.07) higher when running without orthotics (Table 1).

table1

Table 1 Knee loads as a function of orthotic intervention. Notes: * = significant difference p<0.008.

Discussion

This study aimed to determine whether foot orthoses reduce the loads experienced by the patellofemoral joint during the stance phase of running. Previous analyses have examined the efficacy of orthotic devices in the treatment of patellofemoral disorders, but this represents the first investigation to examine the effects of orthotic devices on the loads experienced by the joint itself.

In support of our hypothesis, the key observation from the current investigation is that patellofemoral load parameters were significantly reduced with the presence of orthotic intervention when compared to running without orthotic inserts. This finding may have relevance clinically and serve to provide further insight into the mechanisms by which foot orthoses serve to attenuate the symptoms of patellofemoral pain Ho et al. [19]. The aetiology and pathogenesis of patellofemoral disorders are a function of habitual and excessive loads experienced by the patellofemoral joint itself, which could account for the high incidence of patellofemoral disorders in runners. This current investigation shows that using foot orthoses may be a potential mechanism by which runners are able to attenuate their risk of injury through reductions in knee joint loading.

It is hypothesized that the reductions in patellofemoral kinetics observed in the current study are linked to the additional midsole cushioning associated with the orthotic device. When running with increased midsole cushioning runners typically utilize reduced knee flexion angle at footstrike and throughout the stance phase (Figure 1a). Reductions in knee flexion are associated with lengthening of the quadriceps moment arm, which serves to reduce the load experienced by the patellofemoral joint as PTFC and PTCP are governed by the force generated in the quadriceps [19].

In conclusion, the findings from the current study show that running with foot orthotics are associated with significant reductions in patellofemoral loading parameters when compared to running without orthotic intervention. Given the proposed relationship between the magnitude of patellofemoral loading and the aetiology of patellofemoral pathology, it is proposed that the risk of the developing running related injuries at the patellofemoral joint may be attenuated as a function of orthotic intervention.

Acknowledgements

The authors wish to thank Robert Graydon for his technical assistance during data collection.

References

  1. Denvir MA, Gray GA. Run for your life: exercise, oxidative stress and the ageing endothelium. Journal of Physiology 2009 Sep;587(Pt17):4137-4138. PubMed
  2. Hreljac A. Impact and overuse injuries in runners. Medicine & Science in Sports & Exercise 2004 May;36(5):845-849. PubMed
  3. Marti B, Vader JP, Minder CE, Abelin T. On the epidemiology of running injuries The 1984 Bern Grand-Prix study. American Journal of Sports Medicine 1988 May-Jun;16(3): 285-294. PubMed
  4. Besier TF, Gold GE, Beaupre GS, Delp SL. A modeling framework to estimate patellofemoral joint cartilage stress in vivo. Medicine & Science in Sports & Exercise 2005 Nov;37(11):1924–1931. PubMed
  5. Crossley KM. Is patellofemoral osteoarthritis a common sequela of patellofemoral pain?. British Journal of Sports Medicine 2014 Mar;48(6):409-410. PubMed
  6. Thomas MJ, Wood L, Selfe J, Peat G. Anterior knee pain in younger adults as a precursor to subsequent patellofemoral osteoarthritis: a systematic review. BMC Musculoskeletal Disorders 2010 Sep;11: 201. PubMed
  7. Collins N, Crossley K, Beller E, Darnell R, McPoil T, Vicenzino B. Foot orthoses and physiotherapy in the treatment of patellofemoral pain syndrome: randomised clinical trial. British Medical Journal 2008 Oct; 337:1735. link
  8. Eng JJ, Pierrynowski MR. Evaluation of soft foot orthotics in the treatment of patellofemoral pain syndrome. Physical Therapy 1993 Feb;73(2):62-68. PubMed
  9. Barton CJ, Menz HB, Crossley KM. Clinical predictors of foot orthoses efficacy in individuals with patellofemoral pain. Medicine & Science in Sports & Exercise 2011 Sep;43(9):1603-1610. PubMed
  10. Pitman D, Jack D. A clinical investigation to determine the effectiveness of biomechanical foot orthoses as initial treatment for patellofemoral pain syndrome. Journal of Prosthetics & Orthotics 2000;12(4):110–116. link
  11. Sinclair J, Hobbs SJ, Taylor PJ, Currigan G, Greenhalgh A. The influence of different force measuring transducers on lower extremity kinematics. Journal of Applied Biomechanics 2014 Jul; 40(3):476-479. PubMed
  12. Sinclair J, Edmundson CJ, Brooks D, Hobbs SJ. Evaluation of kinematic methods of identifying gait Events during running. International Journal of Sport Science & Engineering 2011 Aug; 5(3): 188-192. link
  13. 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 Jun;10(4):171-178. PubMed
  14. Ward SR, Powers CM. The influence of patella alta on patellofemoral joint stress during normal and fast walking. Clinical Biomechanics 2004 Dec;19(10):1040–1047. PubMed
  15. Bonacci J, Vicenzino B, Spratford W, Collins P. Take your shoes off to reduce patellofemoral joint stress during running. British Journal of Sports Medicine 2014 Mar;48(6):425-428. PubMed
  16. Kulmala JP, Avela J, Pasanen K, Parkkari J. Forefoot strikers exhibit lower running-induced knee loading than rearfoot strikers. Medicine & Science in Sports & Exercise 2013 Dec;45(12):2306-2313. PubMed
  17. Sinclair J. Effects of barefoot and barefoot inspired footwear on knee and ankle loading during running. Clinical Biomechanics 2014 Apr;29(4):395-399. PubMed
  18. Keino BJ, Powers CM. Patellofemoral stress during walking in persons with and without patellofemoral pain. Medicine & Science in Sports & Exercise 2002 Oct;34(10):1582–1593. PubMed
  19. Ho KY, Blanchette MG, Powers CM. The influence of heel height on patellofemoral joint kinetics during walking. Gait & Posture 2012 Jun;36(2):271-275. PubMed
  20. van Eijden TM, Kouwenhoven E, Verburg J, Weijs WA. A mathematical model of the patellofemoral joint. Journal of Biomechanics 1986;19(3):219–229. PubMed
  21. Powers CM, Lilley JC, Lee TQ. The effects of axial and multiplane loading of the extensor mechanism on the patellofemoral joint. Clinical Biomechanics 1998 Dec;13(8):616–624. PubMed

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)

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.

References

  1. Hreljac A. Impact and overuse injuries in runners. Med Sci Sports Exerc. 2004;36 (5): 845-9. – Pubmed
  2. Van gent RN, Siem D, Van middelkoop M et-al. Incidence and determinants of lower extremity running injuries in long distance runners: a systematic review. Br J Sports Med. 2007;41 (8): 469-80. – Pubmed
  3. Sinclair J, Richards J, Taylor PJ et-al. Three-dimensional kinematic comparison of treadmill and overground running. Sports Biomech. 2013;12 (3): 272-82. – Pubmed
  4. Milgrom C, Finestone A, Segev S et-al. Are overground or treadmill runners more likely to sustain tibial stress fracture? Br J Sports Med. 2003;37 (2): 160-3. –Pubmed
  5. Clement DB, Taunton JE, Smart GW et-al. A survey of overuse running injuries. Phys Sport Med 1981 9: 47-58.
  6. Segesser B, Nigg BM, Morell F. Achillodynia and tibial insertion tendinoses. Med U Sport 1980 20: 79-83.
  7. Sinclair J, Hobbs SJ, Currigan G et-al. Tibiocalcaneal kinematics during barefoot and in barefoot-inspired shoes in comparison to conventional running footwear. Mov Sport Sci 2014 83: 67-75. – link
  8. Nigg BM, Cole GK, Nachbauer W. Effects of arch height of the foot on angular motion of the lower extremities in running. J Biomech. 1993;26 (8): 909-16. – Pubmed
  9. Stergiou N, Bates BT, James SL. Asynchrony between subtalar and knee joint function during running. Med Sci Sports Exerc. 1999;31 (11): 1645-55.  – Pubmed
  10. McClay IS, Manal KT. Coupling parameters in runners with normal and excessive pronation. J App Biomech 1997 13: 109-124. – link
  11. Hintermann B, Nigg BM. The movement transfer between foot and leg in vitro. Sportverletzungen 1994 8: 60-66.
  12. Cappozzo A, Catani F, Croce UD et-al. Position and orientation in space of bones during movement: anatomical frame definition and determination. Clin Biomech (Bristol, Avon). 1995;10 (4): 171-178. – Pubmed
  13. Sinclair J, Taylor PJ, Hobbs SJ. Digital Filtering of Three-Dimensional Lower Extremity Kinematics: an Assessment. J Human Kin 2013 39: 35-37. – link
  14. Sinclair J, Taylor PJ, Edmundson CJ, Brooks D, Hobbs SJ. Influence of the helical and six available Cardan sequences on 3D ankle joint kinematic parameters. Sports Biomech 2012 11: 430–437 – link
  15. Deleo AT, Dierks TA, Ferber R et-al. Lower extremity joint coupling during running: a current update. Clin Biomech (Bristol, Avon). 2004;19 (10): 983-91. – Pubmed
  16. Willems TM, De clercq D, Delbaere K et-al. A prospective study of gait related risk factors for exercise-related lower leg pain. Gait Posture. 2006;23 (1): 91-8. – Pubmed
  17. Lee SY, Hertel J, Lee SC. Rearfoot eversion has indirect effects on plantar fascia tension by changing the amount of arch collapse. Foot (Edinb). 20 (2-3): 64-70.  – Pubmed
  18. Taunton JE, Ryan MB, Clement DB. A prospective study of running injuries: the Vancouver Sun Run “In training” clinics. B J Sports Med 2003 37: 239-244. – link
  19. Duffey MJ, Martin DF, Cannon DW et-al. Etiologic factors associated with anterior knee pain in distance runners. Med Sci Sports Exerc. 2000;32 (11): 1825-32. – Pubmed
  20. Nawoczenski DA, Saltzman CL, Cook TM. The effect of foot structure on the three-dimensional kinematic coupling behavior of the leg and rear foot. Phys Ther. 1998;78 (4): 404-16. – Pubmed
  21. Williams DS, McLay IS, Hamill J et-al. Lower extremity kinematic and kinetic differences in runners with high and low arches. J  App Biomech 2001 17: 153–163. – link
  22. Sinclair J, Greenhalgh A, Edmundson CJ, Brooks D, Hobbs SJ. Gender Differences in the Kinetics and Kinematics of Distance Running: Implications for Footwear Design. Int J Sports Sci Eng 2012 6: 118–128. – link
  23. Stacoff A, Kälin X, Stüssi E. The effects of shoes on the torsion and rearfoot motion in running. Med Sci Sports Exerc. 1991;23 (4): 482-90. – Pubmed
  24. 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 14: 64– 69. – link

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

  1. Sinclair J, Greenhalgh A, Taylor PJ et-al. Differences in tibiocalcaneal kinematics measured with skin and shoe-mounted markers. Human Movement 2013 14: 64– 69. – link
  2. Leardini A, Benedetti MG, Berti L et-al. Rear-foot, mid-foot and fore-foot motion during the stance phase of gait. Gait Posture. 2007;25 (3): 453-62. – Pubmed
  3. Stacoff A, Kälin X, Stüssi E. The effects of shoes on the torsion and rearfoot motion in running. Med Sci Sports Exerc. 1991;23 (4): 482-90. – Pubmed
  4. Stacoff A, Nigg BM, Reinschmidt C et-al. Tibiocalcaneal kinematics of barefoot versus shod running. J Biomech. 2000;33 (11): 1387-95. – Pubmed
  5. Bishop C, Thewlis D, Uden H et-al. A radiological method to determine the accuracy of motion capture marker placement on palpable anatomical landmarks through a shoe. Footwear Sci 2011 3: 169–177. – link
  6. Bishop C, Paul G, Thewlis D. The development of a kinematic model to quantify in-shoe foot motion. J Foot Ankle Res2012 5: S43. – link
  7. Reinschmidt C, Stacoff A, Stüssi E. Heel movement within a court shoe. Med Sci Sports Exerc. 1992;24 (12): 1390-5. – Pubmed
  8. Sinclair J, Hobbs SJ, Taylor PJ et-al. The influence of different force and pressure measuring transducers on lower extremity kinematics measured during running. J Appl Biomech. 2014;30 (1): 166-72. – Pubmed
  9. Sinclair J, Edmundson CJ, Brooks D et-al. Evaluation of kinematic methods of identifying gait events during running. Int J Sports Sci Eng 2011 5: 188–192. – link
  10. Cappozzo A, Catani F, Croce UD et-al. Position and orientation in space of bones during movement: anatomical frame definition and determination. Clin Biomech (Bristol, Avon). 1995;10 (4): 171-178. – Pubmed
  11. Shultz R, Jenkyn T. Determining the maximum diameter for holes in the shoe without compromising shoe integrity when using a multi-segment foot model. Med Eng Phys. 2012;34 (1): 118-22. – Pubmed
  12. Winter DA. Biomechanics and motor control of human movement. John Wiley and Sons, Inc., New York, 1990.
  13. Ferrari A, Cutti AG, Cappello A. A new formulation of the coefficient of multiple correlation to assess the similarity of waveforms measured synchronously by different motion analysis protocols. Gait Posture. 2010;31 (4): 540-2. – Pubmed
  14. Carson MC, Harrington ME, Thompson N et-al. Kinematic analysis of a multi-segment foot model for research and clinical applications: a repeatability analysis. J Biomech. 2001;34 (10): 1299-307. – Pubmed
  15. Buchanan KR, Davis I. The relationship between forefoot, midfoot, and rearfoot static alignment in pain-free individuals. J Orthop Sports Phys Ther. 2005;35 (9): 559-66. –Pubmed
  16. Nishiwaki T, Nakaya S. Footwear sole stiffness evaluation method corresponding to gait patterns based on eigenvibration analysis, Footwear Science, 2009: 95-101.- Link
  17. Pohl MB, Messenger N, Buckley JG. Forefoot, rearfoot and shank coupling: effect of variations in speed and mode of gait. Gait Posture. 2007;25 (2): 295-302. – Pubmed
  18. Powers CM. The influence of altered lower-extremity kinematics on patellofemoral joint dysfunction: a theoretical perspective. J Orthop Sports Phys Ther. 2003;33 (11): 639-46. – Pubmed
  19. Stergiou N, Bates BT, James SL. Asynchrony between subtalar and knee joint function during running. Med Sci Sports Exerc. 1999;31 (11): 1645-55. – Pubmed
  20. Tiberio D. Evaluation of functional ankle dorsiflexion using subtalar neutral position. A clinical report. Phys Ther. 1987;67 (6): 955-7. – Pubmed

 

 

A Review of the Function of the Quadratus Plantae

by Thomas P. Lyman, B.S.1  

The Foot and Ankle Online Journal 2 (11): 5

There has been very little scientific evidence of any well defined function of the quadrates plantae (QP). It has been the focal point of multiple theories. This paper discusses four of those theories: the QP as a counter to the oblique force of the flexor digitorum longus (FDL), QP as a pronator of the foot, the QP as a plantarflexor of the lesser digits and lastly the QP as a stabilizer of the lumbrical muscles. No single theory has received unanimous support. The following is a presentation of the support and opposition which has surfaced for each theory.

Key Words: Quadratus plantae, plantar foot muscles, biomechanics.

This is an Open Access article distributed under the terms of the Creative Commons Attribution License.  It permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. ©The Foot and Ankle Online Journal (www.faoj.org)

Accepted: October, 2009
Published: November, 2009

ISSN 1941-6806
doi: 10.3827/faoj.2009.0211.0005


The quadratus plantae (QP) muscle is located in the second layer of the plantar musculature and consists of two heads separated by the long plantar ligament. (Fig. 1)

Figure 1*  The qudratus plantae muscle resides in the second muscle layer of the foot.  It’s origin is from the calcaneus and inserts at the origin of the lumbrical muscles.    

 *This faithful reproduction of a lithograph plate from Gray’s Anatomy, a two-dimensional work of art, is not copyrightable in the U.S. as per Bridgeman Art Library v. Corel Corp.; the same is also true in many other countries, including Germany. Unless stated otherwise, it is from the 20th U.S. edition of Gray’s Anatomy of the Human Body, originally published in 1918 and therefore lapsed into the public domain. 

The medial head originates from the medial surface of the plantar calcaneus, while the lateral head originates from the lateral border of the calcaneus. The muscle inserts into the posterolateral margin of the flexor digitorum longus (FDL) tendon as it crosses the plantar foot obliquely from posteromedial to anterolateral. This insertion falls at the level of the origin of the lumbrical muscles. [1] As for the biomechanical function of the QP, there is a relative lack of evidence supporting any one particular theory. At least four theories of its function have been proposed, however very little has been demonstrated through empirical research. The goal of this paper is to address these four theories, along with a discussion of the consenting and dissenting voices of each.

The QP Countering the Oblique Pull of the FDL

Certainly, one of the most widely accepted theories states that the QP functions to balance out the oblique pull of the FDL on the lesser digits. Again, the FDL tendon courses obliquely across the plantar foot in a direction that would tend to pull the lesser digits, particularly the fourth and fifth digits, into an adductovarus position, if it were not for this straightening effect. [1-8] Although many have simply accepted this theory and published it as fact, others have tried to provide direct evidence. Manoli and Weber reported on three cases which provided fairly convincing evidence in support of this theory. [9]

Their intention was to demonstrate that the QP was in its own separate plantar compartment, the calcaneal compartment. Each of the cases consisted of a calcaneal fracture which healed uneventfully, but with lesser digit adductovarus contractures developing in each of the patients within 8-13 months. They also found that none of the deformities were reducible with ankle plantarflexion, indicating intrinsic muscular contracture. They hypothesized that the contractures were due to an undetected calcaneal compartment compression damaging the neurovascular supply to the QP. They concluded therefore that the QP functions to prevent adductovarus toe contractures of the lesser digits.

Interestingly however, the effect of the damaged QP in this study was most prominent in the second and third digits. If the QP was damaged and the oblique angle of the FDL was exaggerated on the toes, wouldn’t the fourth and fifth receive the most oblique force? The results didn’t appear completely consistent with the theory.

Not everyone has accepted this theory. Basing much of their argument on the anatomic presence of the flexor tendon sheath, others claim that an oblique pull on the toes would be impossible. [10,11] The well-defined flexor sheath on the plantar side of each digit begins proximal to the metatarsal phalangeal joint. The tendon of the FDL runs in this sheath and its resultant force could only be in line with the sheath. Furthermore, the interphalangeal joints are hinge joints and any force applied to them could only result in uniplanar motion. [10,11] Kaplan took his argument one step further with a cadaver experiment. He applied multidirectional forces to the cadaver FDLs and the result in all cases was straight toe flexion. He did not find that the QP was required to produce straight flexion of the digits. [10]

The QP as a Pronator

This experiment conducted by Kaplan contributed to the development of the theory that the primary function of the QP is to pronate the foot. In his cadaver experiments, the FDL was pulled on both a fixed calcaneus as well as an unfixed calcaneus. When the FDL was pulled on the unfixed calcaneus, it resulted in straight toe flexion with obvious adduction of the foot and supination of the lateral border. When the calcaneus was unfixed and a force was applied to the FDL and QP, the resultant motion was again straight toe flexion, but this time the foot pronated with no adduction. Therefore the QP seemed to be a pronator of the foot. Kaplan also dissected the feet of two infant cadavers which were born with congenital club foot. In further support of the theory, there were no QP muscles present in either one. [10]

Reeser, et al., challenged this theory with electromyographic studies of the foot. The study revealed no difference in the activity of the QP when the foot was inverted versus everted. Their conclusion was that the QP does not necessarily aid more in pronation or supination. They theorized that a pronator muscle would fire more during pronation than supination, which apparently was not the case. This study was based on four participants. [11]

The QP as a Lesser Digit Plantarflexor

Reeser, et al.,’s study brought up a point in support of yet another theory. They found that the only time that the QP consistently fired in each of the patients was during plantarflexion of the digits. The other movements incorporated with the toe plantar flexion did have some impact on the strength of the contraction, but ultimately combinations of motion which included plantarflexory toe movements demonstrated firing of the QP.

They concluded that this muscle was intimately involved with toe plantarflexion. They proposed that the QP functions to counteract active insufficiency of the FDL in plantarflexion, as well as create more force at the interphalangeal joints when needed. [11] The previously discussed case series presented three cases of supposed QP damage followed by claw toes of the second and third digits. [9] A contracture of the QP may have been the force pulling the distal phalanges into plantarflexion. It was interesting however that the fourth and fifth digits were not the primary digits affected.

The QP as a Lumbrical Stabilizer

Anatomic location is the basis for the next theory proposed by Root, et al. [2] They stated that due to the location of insertion of QP on the FDL, it could possibly stabilize the lumbricales. The lumbricales originate on the FDL near the insertion of the QP. [2] This theory has not met with the same opposition that the others have. This probably provides a small piece of the complete puzzle, whatever that may be.

Conclusion

No single theory about the function of the QP has been unanimously supported or rejected by the scientific community. There is indeed a lack of evidence available to support any one line of thinking. The reality of this controversy is that a combination of these theories, and some yet to be discovered, make up the true function of the QP muscle. Certainly more research is warranted. If the first theory discussed holds true, and a weak QP causes adductovarus deformity of the fourth and fifth digits, then a method of strengthening the QP could be developed to prevent such cases. Likewise the involvement of the QP in the other functions discussed could lead to advances in the prevention and correction of some supinatory deformities and contracted digits.

Insight into the etiology of this deformity, and others like it, can lead to prevention. The impact of the QP needs to be determined and its function isolated to lead to these important steps of foot control and function.

References

1. Moore K, Dalley A: Clinically oriented anatomy. Fifth edition. pp 555 – 725, Lippincott, Williams and Wilkins, Philadelphia, 2006.
2. Root ML, Orien WP, Weed JH: Functions of the muscles of the foot. In Normal and Abnormal Function of the Foot, pp250 – 252, edited by ML Root, WP Orien, JH Weed, Clinical Biomechanics Corporation, Los Angeles, 1977.
3. Jones FW: Structure and function as seen in the foot. p 166, London (UK): Balliere, Tindall and Cox; 1944.
4. Lewis OJ: The comparative morphology of M. flexor accessories and the associated long flexor tendons. J Anat 96: 321 – 333, 1962.
5. Sgarlato TE: Pathomechanics of various developmental abnormalities. In A Compendium of Podiatric Biomechanics, pp 369 – 422, edited by TE Sgarlato, California College of Podiatric Medicine, San Francisco, 1971.
6. McGlamry ED, Jimenez AL, Green DR: Deformities of the intermediate digits and the metatarsophalangeal joint. In McGlamry’s Comprehensive Textbook of Foot and Ankle Surgery, pp 253 – 304, edited by AS Banks, MS Downey, DE Martin, SJ Miller, Lippincott, Williams and Wilkins, Philadelphia, 2001.
7. Morris JL: Biomechanical implications of hammertoe deformities. Clin Podiatr Med Surg 3 (2): 339 – 346, 1986.
8. Sooriakumaran P, Sivananthan S: Why does man have a quadratus plantae? A review of its comparative anatomy. Croat Med J 46(1):30-35, 2005.
9. Manoli A, Weber TG: Fasciotomy of the foot: An anatomical study with special reference to release of the calcaneal compartment. Foot Ankle 10: 267 – 275, 1990.
10. Kaplan EB: Morphology and function of the quadratus plantae. Bull Hosp Joint Dis 20: 84 – 95, 1959.
11. Reeser LA, Susman RL, Stern JT: Electromyographic studies of the human foot: experimental approaches to hominid evolution. Foot Ankle 3: 391 – 407, 1983.


Address correspondence to: Thomas P. Lyman, 2639 Wentworth Rd, Philadelphia, PA 19131
Email: tub09643@temple.edu

Temple University School of Podiatric Medicine, Student.

© The Foot and Ankle Online Journal, 2009