JCLB-03812; No of Pages 7 Clinical Biomechanics xxx (2014) xxx–xxx
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Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running Gail M. Monaghan a,⁎, Wen-Hao Hsu a, Cara L. Lewis a, Elliot Saltzman a, Joseph Hamill b, Kenneth G. Holt a a b
Department of Physical Therapy & Athletic Training, College of Health and Rehabilitation Sciences: Sargent College, Boston, MA, USA Department of Kinesiology, University of Massachusetts, Amherst, MA, USA.
a r t i c l e
i n f o
Article history: Received 25 September 2013 Accepted 19 June 2014 Keywords: Pronation Forefoot inversion Rearfoot inversion Musculoskeletal injury
a b s t r a c t Background: Clinically, foot structures are assessed intrinsically – relation of forefoot to rearfoot and rearfoot to leg. We have argued that, from a biomechanical perspective, the interaction of the foot with the ground may influence forces and torques that are propagated through the lower extremity. We proposed that a more appropriate measure is an extrinsic one that may predict the angle the foot makes with ground at contact. The purposes of this study were to determine if the proposed measure predicts contact angles of the forefoot and rearfoot and assess if the magnitude of those angles influences amplitude and duration of foot eversion during running. Methods: With the individual in prone, extrinsic clinical forefoot and rearfoot angles were measured relative to the caudal edge of the examination table. Participants ran over ground while frontal plane forefoot and rearfoot contact angles, forefoot and rearfoot eversion amplitude and duration were measured. Participants were grouped twice, once based on forefoot contact inversion angle (moderate b median and large N median) and once based on rearfoot contact inversion angle (moderate b median and large N median). Findings: The forefoot and rearfoot extrinsic clinical angles predicted, respectively, the forefoot and rearfoot angles at ground contact. Large forefoot contact angles were associated with greater amplitudes (but not durations) of forefoot and rearfoot eversion during stance. Rearfoot contact angles, however, were associated with neither amplitudes nor durations of forefoot and rearfoot eversion. Interpretation: Possible mechanisms for the increased risk of running injuries associated with large forefoot angles are discussed. © 2014 Elsevier Ltd. All rights reserved.
1. Introduction Thirty million Americans run to maintain health and fitness (Surgeon General, 1999). The typical runner has become more competitive, entering half-marathon races and logging in 1269 miles/year (Lamppa, 2011). One half of runners will experience musculoskeletal injuries which will stop them running for at least a week (Buist et al., 2010; Jacobs and Berson, 1986) and the cumulative effect of these injuries over many years of running may be knee and hip osteoarthritis (Brandt et al., 2008). As the foot contacts the ground many times, with ground reaction forces up to three times body weight, it is expected that foot structure will affect the stresses experienced not only by tissues reported as common running injury sites (i.e., patellofemoral joint, the iliotibial band and knee cartilage; (Taunton et al., 2002)), but also by the hip joint. Despite the prevalence of musculoskeletal injury
⁎ Corresponding author at: Boston University, College of Health and Rehabilitation Sciences: Sargent College, Department of Physical Therapy and Athletic Training, 635 Commonwealth Avenue, Boston, MA 02215, USA. E-mail address: [email protected] (G.M. Monaghan).
associated with running, the biomechanical mechanisms linking foot structure to abnormal movement patterns and potential injury are poorly understood (Razeghi and Batt, 2002). A critical factor that may contribute to this lack of understanding may be the way that foot structure is defined. Traditionally, foot structure is defined intrinsically, that is, relative to the other adjacent segments (forefoot to rearfoot, rearfoot to leg) (Root et al., 1977) (Fig. 1: Intrinsic measure). However, the intrinsic measure may not adequately represent the angle of the foot as it contacts the ground and, therefore, may not adequately capture the effects that contact angle may have on the forces and torques generated around the foot and upward through the kinematic chain. Additionally, given its poor inter-rater reliability (Van Gheluwe et al., 2002), using the intrinsic clinical measure of foot structure may contribute to the poor correlation of this measure to foot kinematics observed during the stance phase of locomotion (Hamill et al., 1989). Holt and Hamill (1995) proposed that the angle between the foot and the ground at foot contact during locomotion is the critical determinant of the force and torque patterns created, first at the foot and then through the rest of the body. In the current study of running, with the individual prone, clinical measures of rearfoot and forefoot angles
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Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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G.M. Monaghan et al. / Clinical Biomechanics xxx (2014) xxx–xxx
Fig. 1. Clinical foot measures. Intrinsic measure of clinical foot angle (left): rearfoot (B) and forefoot (A) angles are measured relative their respective proximal segments. The rearfoot angle is measured as the angle between the bisection of the calcaneus (black solid line) and the bisection of the lower leg (dashed black line). The forefoot angle is measured as the angle between the line through the metatarsal heads (black solid line) and the line bisecting the calcaneus. Extrinsic measure of clinical foot angle (right): rearfoot (C) and forefoot (D) angles are both measured relative to the caudal edge of the table. Rearfoot angle is measured as the angle between a line bisecting the calcaneus (white solid line) and a line perpendicular to the caudal edge of the table which is parallel to the longitudinal axis of the table (dashed black line). Forefoot angle is measured as the angle between the line connecting the first through fifth metatarsal heads (black solid line) and the caudal edge of the table (black dashed line).
were assessed extrinsically relative to a fixed environmental axis, the caudal edge of the table. This edge provides an environmental reference line that is functionally meaningful for locomotion in the sense that, were the table and subject to be rotated to an upright position, this edge would be parallel to the ground. Using this same method, Monaghan et al. (2013) previously showed that the angle of the forefoot to the ground at forefoot contact determines the amount and duration of eversion during walking, whereas the rearfoot angle at rearfoot contact had no effect on the amplitude or duration of eversion during walking. It is not known, however, if these findings also apply to running gait. Rearfoot kinematics during locomotion in the laboratory have been studied extensively (Davis et al., 2003; McClay and Manal, 1997; Pohl and Buckley, 2008) and have contributed to understanding injury mechanisms, but there is no epidemiological evidence to suggest that rearfoot angles measured clinically are associated with injury. Forefoot kinematics, however, have been less well studied. While some research has tested the application of foot orthoses to address clinically measured forefoot abnormalities (Ferber et al., 2005; Stackhouse et al., 2004; Williams et al., 2003), few studies report forefoot kinematics during locomotion or investigate the association between clinical forefoot abnormality and injury (Gross et al., 2007; Korpelainen et al., 2001; Lun et al., 2004). Holt and Hamill (1995) proposed that a clinical forefoot inversion deformity, measured in a non-weight bearing position, may be a prime contributor to the occurrence of pronation between heel off and toe off, which is referred to as “late” pronation. This late pronation could disrupt the coordinated movement of the foot, knee, and hip, and result in injury of proximal joints, a claim supported by others (Michaud, 1993; Tiberio, 1987). The effects of late forefoot pronation on rearfoot kinematics may be greater during running due to the presence of correspondingly larger ground reaction forces and resultant torques than during walking. Therefore, the purposes of this study were: 1) to test the ability of the extrinsic clinical measures of rearfoot and forefoot angles to predict the segment-to-ground angles of both rearfoot and forefoot at initial ground contact and 2) to explore the relationship between rearfoot and forefoot angles at initial contact and kinematic variables during the stance phase of running. 2. Methods 2.1. Subjects and instrumentation Eight females and six males (mean (SD) age 23.6 (4.9) years) were recruited for this study. The average and standard deviation mass and
height of participants were 66 (10.5) kg and 1.7 (0.1) m respectively. The inclusion criterion for all participants was to have an extrinsically measured forefoot angle greater than 15° of inversion. We selected this range because Gross et al. (2007) suggested that only the extremes of forefoot structural abnormality had an impact on injury. Individuals with current lower extremity pain were excluded. Prior to participation, each participant signed a consent form approved by the Institutional Review Boards of Boston University and University of Massachusetts, Amherst. Photographs of each subject's feet were taken using a six megapixel digital camera, Nikon Coolpix L11 which was fixed to the wall. Each subject was fitted with standardized sandals (Bite (LLC) running sandals with straps at the forefoot and rearfoot). An 8-camera Oqus motion capture system (Qualisys Medical AB, Gothenburg, Sweden) was used to record foot segment motions at 240 Hz. Ground reaction force data from a force plate (Advanced Mechanical Technologies, Inc., Watertown, MA, USA). Force plate data were synchronized with the motion data using Qualisys Track Manager. 2.2. Data collection On the first of three visits to the laboratory, the participant was screened to ensure that inclusion criteria were met. On the second visit, the participant received a clinical evaluation while lying in a prone position on an examination table. Briefly (see Monaghan et al., 2013), the medial malleoli were manually aligned at the caudal edge of the table, the legs were manually placed with the medial malleoli 33 cm apart and in neutral rotation, and, using light pressure at the third metatarsal head the examiner guided the ankle into 0° of dorsiflexion. The participant was instructed to hold the position after the examiner removed pressure from the third metatarsal head. A digital photograph was then taken. This positioning process was performed three times for each participant, and three digital pictures were obtained. Participants practiced running along a 10 m runway to ensure that the right foot contacted a force plate embedded in the runway. The participant's preferred speed was determined as the average speed from three such runs. In order to test the effect of small changes in running speed, the participant was then asked to run at slower and faster speeds than the preferred speed until consistent fast and slow speeds were reached. It was required that speed levels differ by at least 0.2 m/s. Only trials within approximately 0.1 m/s of the designated preferred, slow and fast speeds were accepted. Running speed was monitored using a set of photocell-triggered timers. Eleven retro-reflective
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
G.M. Monaghan et al. / Clinical Biomechanics xxx (2014) xxx–xxx
markers (9.5 mm in diameter) were placed on the foot and ankle according to the Leardini et al.'s model (2007), and a 12th ‘virtual marker’ was defined computationally as the midpoint between the sustentaculum tali and the peroneal tubercle markers (Fig. 2). Following an upright standing calibration trial, each participant performed five acceptable running trials at each of the three speeds. Approximately 1 month later on visit three, the motion capture procedures were repeated but without a clinical evaluation. 2.3. Data processing Rearfoot and forefoot extrinsic clinical angles were measured from the digital photographs using commercially available software (CanvasX, ACD Systems, Vancouver, British Columbia, Canada). The detailed protocol is described by Monaghan et al. (2013). Referencing the clinical angle to the environment (caudal edge of the table), rearfoot angle was defined as the angle between a line bisecting the calcaneus and a line perpendicular to the caudal edge of the table (Fig. 1C); forefoot angle was defined as the angle between the line connecting the first and fifth metatarsal heads and the caudal edge of the table (Fig. 1D). This edge defines a line that is parallel to the mediolateral axis in the participant's frontal
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plane and is parallel to the ground when standing. The average of measurements taken from the three photographs was used as the clinical extrinsic foot angle. The ICC's for intra-rater reliabilities were 0.91 for the forefoot and 0.87 for the rearfoot. Gross et al. (2007) reported similar intra-rater reliabilities, and also reported inter-rater reliabilities of 0.93 for the forefoot and 0.85 for the rearfoot. Marker data were labeled using Qualisys Track Manager. Raw kinematic data were low-pass filtered using a fourth order, zero-lag Butterworth digital filter with a cut-off frequency of 8 Hz. Three dimensional segment angles were calculated in Visual 3D using an x (flexion/extension), y (eversion/inversion), and z (axial rotation) Cardan rotation sequence (Cole et al., 1993). Eversion/inversion segment angles of the rearfoot and forefoot were calculated in the frontal plane relative to the horizontal and were used as proxies for pronation/supination. Using segment angles instead of joint angles assesses the foot motion relative to the ground. Time derivatives of the angular position time series were calculated in Visual 3D. Initial contact and toe-off were determined using the vertical ground reaction force. Stance phase data were normalized to percent stance duration. All runners were rearfoot strikers. Rearfoot contact was defined as initial contact; forefoot contact was defined when the third derivative (jerk) of the vertical position data of the fifth metatarsal head marker (VMH in Fig. 2) equaled zero, and heel off was assumed when the jerk of the vertical position data of the lateral heel marker (PT in Fig. 2) equaled zero (Winter, 2005). The onsets of rearfoot and forefoot inversion were defined when the velocities of the rearfoot and forefoot angles, respectively, in the frontal plane were greater than zero and remained above zero for the remainder of stance. Time series data for the segment angles were exported from Visual 3D. Custom Matlab (Mathworks, Natick, MA) programs extracted rearfoot and forefoot angle at foot contact, amplitude of rearfoot and forefoot eversion, duration of rearfoot and forefoot eversion, and duration (% stance) between heel off and the onset of rearfoot and forefoot inversion. Amplitude of rearfoot eversion was defined as the difference between the rearfoot angle at contact and its maximum eversion angle. Amplitude reflects the motion during the loading phase instead of focusing on a single time point. Duration of rearfoot eversion was defined as the percent of stance between contact and the onset of rearfoot inversion and therefore includes the time when the foot is neither inverting nor everting. Duration of forefoot eversion was defined similarly as the percent of stance between contact and the onset of forefoot inversion. Durations between heel off and the respective onsets of rearfoot and forefoot inversion were measured as percent of stance. 2.4. Participant categorization
Fig. 2. Foot markers, landmarks, and segments. In accordance with the model of Leardini et al. (2007) retro-reflective markers (solid circles) were placed on the bony landmarks of the foot: navicular (TN), first (FMB), second (SMB), and fifth (VMB) metatarsal bases, first (FMH), second (SMH) and fifth (VMH) metatarsal heads and the most distal and dorsal point of the head of the hallux (PM). Markers on the calcaneus were located at the triceps surae insertion (CA), the most medial apex of the sustentaculum tali (ST) and the lateral apex of the peroneal tubercle (PT). IC (gray circle) is intermedius calcaneus, an additional computed landmark (‘virtual marker’) defined at the midpoint between ST and PT. The rearfoot segment and forefoot segments are red dashed-line triangles. The rearfoot segment was defined by the triangle CA–ST–PT; its origin is located at CA, and its x-axis is defined along the line from the origin (CA) to IC. The forefoot segment was defined by the triangle SMB–FMH–VMH; its origin is located at SMB, and its x-axis is defined along the projection of the line from SMB to SMH onto the transverse plane passing through the origin (SMB), FMH and VMH. Image adapted from www.fotosearch.com. k8433317 “Bones of the Human Foot”. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
Participants in this study were grouped twice, once based on their forefoot contact angle and once based on the rearfoot contact angle. The groups were defined by dividing the participants into two equal groups based on the median contact angle: those participants above the median contact angle were in the “large inversion angle” group, and those below the median contact angle were in the “moderate inversion angle” group. This dual categorization allowed us to investigate the influence of contact angle on kinematic parameters separately for forefoot and rearfoot. We chose to compare moderate and large inversion angles, as only participants with very large forefoot inversion angles have been shown to be at risk of injury (Gross et al., 2007). This procedure allowed us to examine the relationship of forefoot and rearfoot contact angles with other foot-related kinematic measures obtained during the stance phase of running. 2.5. Statistical analysis A linear general estimating equation (GEE) analysis was used to determine if mean forefoot and rearfoot extrinsic clinical angles could predict mean forefoot and rearfoot angles at forefoot and rearfoot contact,
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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respectively, during running. Within-subject variables were speed (preferred, slow, fast) and visit (day 1 and day 2). To test if forefoot and rearfoot angles at contact predicted foot kinematics during stance, two separate three factor ANOVAs were performed – one for the groups based on forefoot angle at contact and one for the groups based on rearfoot angle at contact – in order to determine if there was a significant difference between groups in their predictions of foot kinematics during stance. The three factors were 1) group based on the magnitude of the contact angle (large and moderate), 2) running speed (preferred, slow, fast) and 3) visit (day 1 and day 2). The dependent variables were: 1) the amplitude of rearfoot eversion, 2) the amplitude of forefoot eversion, 3) the duration of rearfoot eversion (% stance), 4) the duration (% stance) of forefoot eversion, 5) the duration (% stance) between heel off and the onset of rearfoot inversion, and 6) the duration (% stance) between heel off and the onset of forefoot inversion. We were interested in the main effect of group and any interaction effect between group and speed or visit. Statistical significance was set at an alpha level of 0.05.
Table 1 Characteristics of groups based on contact angle (mean (SD)).
3. Results
moderate rearfoot inversion contact angle groups (Table 2, bottom) (Fig. 3B).
3.1. Effect of speed and visit There were no interaction effects between group and speed or group and visit. Therefore, we focused our analyses on the main effect of group. The average slow, preferred and fast speeds were 2.97, 3.48 and 4.19 m/s. 3.2. Relationships between clinical and contact angles 3.2.1. GEE analysis The P-value tests whether the clinical angle significantly predicts the contact angle, and the slope (B) can be used to describe how the expected (or mean) contact angle changes as the clinical angle score changes. The Wald statistic is a parametric statistic used to test how well the true value of a parameter can predict another parameter whenever the relationship between two parameters can be expressed as a model. The clinical forefoot angle predicted the forefoot angle at contact during running with a positive linear relation (B = 0.14, W(1, 11) = 5.1, P = 0.02). The clinical rearfoot angle also predicted the rearfoot angle at contact, but with a negative linear relation indicating an inverse relationship (B = −.18, W(1, 7) = 25.4, P ≤ 0.01). There was no effect of running speed for the forefoot contact angle. Rearfoot contact angle increased significantly as the speed increased (mean 3.54° [SD 0.48°], mean 4.55° [SD 0.43°], mean 4.60° [SD 0.51°], P = 0.01). 3.2.2. Group comparisons The group with large forefoot inversion angles at forefoot contact had larger clinical angles (Table 1) than the group with moderate forefoot inversion angles (F(1,12) = 11.7, P = 0.01), reflecting the positive linear relation between forefoot clinical angle and forefoot contact angle reported in the regression analysis. Conversely, the group with large rearfoot inversion contact angles at rearfoot contact had smaller clinical angles than the group with moderate rearfoot inversion contact angles (F(1,8) = 6.7, P = 0.03), reflecting the negative linear relation between rearfoot clinical angle and rearfoot contact angle reported in the regression analysis. 3.3. Relationship of contact angles to foot motion 3.3.1. Amplitude of eversion The large forefoot inversion contact angle group had greater amplitude of forefoot and rearfoot eversion (Table 2) (Fig. 3A) than the moderate forefoot inversion group. There were no differences in amplitudes of forefoot or rearfoot eversions between the large rearfoot and the
Forefoot grouping
Moderate
Large
Forefoot angle at contact Rearfoot angle at contact Age Weight Height Forefoot clinical angle Rearfoot clinical angle
3.2° (1.7°) 3.8° (2.4°) 24.9 (6.5) years 70.1 (13.4) kg 1.7 (0.12) m 19.4° (7.1°) 3.6° (10.1°)
7.0° (1.0°) 5.2° (2.0°) 22.4 (2.7) years 63.1 (9.9) kg 1.7 (0.14) m 28° (2.7°) 5.7° (5.2°)
Rearfoot grouping
Moderate
Large
Rearfoot angle at contact Forefoot angle at contact Age Weight Height Forefoot clinical angle Rearfoot clinical angle
2.9° (1.4°) 4.9° (3.6°) 21 (1.4) years 63.3 (15.1) kg 1.6 (0.12) m 27° (4.3°) 9.9° (7.1°)
6.1° (1.4°) 6.1° (3.1°) 22.4 (2.4) years 68.2 (12.7) kg 1.7 (0.13) m 20.6° (8.3°) −0.9° (6.2°)
3.3.2. Duration of eversion There were no differences between groups based on either forefoot or rearfoot contact angle in the duration of forefoot or rearfoot eversion (Table 2). 3.3.3. Duration of stance between heel off and onset of inversion There were no differences between groups based on either forefoot or rearfoot contact angle in the durations of stance between heel off and the onsets of both forefoot inversion and rearfoot inversion (Table 2). 4. Discussion The first goal of this study was to examine the predictive validity of a hypothesized extrinsically defined clinical foot angle measured during non-weight bearing. It was shown that forefoot angle, relative to a fixed environmental reference plane in non-weight bearing, can predict frontal plane angle of the forefoot at contact when running. This new extrinsic measure of foot angle showed predictive validity and had high intra-rater reliability for the forefoot and the rearfoot. However, the finding that a greater clinical rearfoot inversion angle measured in non-weight bearing predicted a smaller rearfoot angle at contact calls into question the common clinical assumption that the non-weight bearing measure of rearfoot is a critical factor in causing excessive pronation and for designing orthoses. According to Holt and Hamill (1995) and the findings of Lafortune et al. (1994), a large forefoot inversion angle at ground contact can produce large pronation torques that result in increased forefoot and rearfoot eversion. The finding in the current study of greater amplitude of forefoot and rearfoot eversion for individuals with a large clinical forefoot angle supports these claims (Fig. 4). Despite the emphasis on the rearfoot in orthotic prescription and in musculoskeletal injury research, rearfoot angle at contact did not appear to be associated with any of the kinematic variables measured during stance. The amplitude and duration of rearfoot eversion may have been influenced by forefoot amplitude and thus masked the association between the rearfoot contact angles and rearfoot kinematic variables. In this paper, we have shown that the foot orientation at ground contact has an important influence on foot biomechanics. This is a unique finding because it provides evidence to support the hypothesis that foot structure and its biomechanical interaction with the ground is a critical first step when investigating the causes of foot and proximal injuries. To fully support the relationship between foot structure and injury, a second step is to show that the high amplitude and velocity of
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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Table 2 Differences in foot contact groups: mean, SD, F ratios and P values for each dependent variable. Variable
Amplitude of forefoot eversion (°) Amplitude of rearfoot eversion (°) Duration of forefoot eversion (%) Duration of rearfoot eversion (%) Duration between heel off and onset of forefoot inversion (%) Duration between heel off and onset of rearfoot inversion (%)
Group
Effect size
Moderate
Large
FF group
Mean (SD)
Mean (SD)
F(11)
P
ES
7.6 (1.1) 3.7 (0.4) 44.9 (1.5) 52.1 (6.1) 13.9 (2.0) 24.5 (4.0)
11.6 (1.1) 6.1 (0.4) 44.5 (1.5) 65.2 (5.5) 12.1 (2.1) 32.6 (3.6)
7 20 0.04 2.6 0.42 5.6
0.02* b0.01* 0.85 0.15 0.53 0.16
3.6 6 −0.26 0.37 −0.88 2.1
Moderate
Large
RF group effect
Mean (SD)
Mean (SD)
F(8)
10.3 (1.7) 5.0 (0.7) 44.8 (1.6) 58.7 (6.4) 14.0 (2.1) 29.6 (4.1)
10.1 (1.7) 4.8 (0.7) 43.1 (1.7) 60.2 (7.1) 11.0 (2.1) 28.0 (4.6)
Differences in rearfoot contact groups: mean, SD, F ratios and P values for each dependent variable Variable
Amplitude of forefoot eversion (°) Amplitude of rearfoot eversion (°) Duration of forefoot eversion (%) Duration of rearfoot eversion (%) Duration between heel off and onset of forefoot inversion (%) Duration between heel off and onset of rearfoot inversion (%)
Group
segment eversion found in this study are associated with changes in proximal motions and torques. We have previously shown that a forefoot and rearfoot posted foot orthotic reduces the motions and torques at the foot and in the ankle, knee and hip during walking (Hsu et al., accepted for publication, J. Biomechanics). A final step is to show how these stresses result in tissue injuries. While our epidemiological study has shown a 5-times greater occurrence of hip replacements in elderly individuals with large forefoot varus (Gross et al., 2007), a similar study is needed for a large group of runners. In addition to these studies, it will also be important to investigate the interaction between foot orientation, and lower extremity joint angles (including forefoot to rearfoot) at foot contact. For example, in addition to foot orientation, the influence of frontal plane shank position relative to the foot will likely influence the motion and torques at the ankle and leg. When taken together, the foot contact angle and joint angles will present a more complete picture of the causes of non-traumatic lower extremity musculoskeletal injury. The comparison of groups with moderate and large amounts of inversion abnormality provided the most stringent of tests for the effects of structure on the kinematics of the foot during running. Nevertheless, it would be interesting and informative to have included a group with minimal or no inversion deformity to compare our groups to ‘normal’ feet. Furthermore, in this study all subjects had a prior history of some
Effect size
0.01 0.05 0.71 0.03 1.3 0.07
P
ES 0.94 0.83 0.42 0.88 0.28 0.79
−0.12 −0.29 −1 0.22 −1.42 −0.37
form of injury either to the feet or lower extremity. However, we made no attempt to investigate the differential effects of type of foot structure associated with specific injuries, particularly to more proximal structures. In future studies we plan to a) investigate the effects of foot structure on the biomechanics of more proximal segments, and b) explore the reasons that different forms of chronic, non-traumatic injury may be related to similar structural abnormalities. The finding that individuals with a greater forefoot inversion angle at contact also had greater forefoot and rearfoot eversion amplitude but not greater eversion duration suggests a mechanism for the occurrence of stresses in distal or proximal regions of the lower extremity. If greater pronation amplitude occurs in the same duration of pronation as individuals with a lesser forefoot inversion angle at contact, then individuals with a greater inversion contact angle had greater forefoot and rearfoot eversion velocities. Subsequently there will be greater tibial and femoral internal rotation velocity and amplitude. The foot of a runner contacts the ground many times each year. For a runner with a large forefoot inversion angle muscles which slow rearfoot eversion and femoral and tibial rotation will have to work harder to control faster movement of greater amplitude. The greater torque generated by these muscles may result in soft tissue injuries (Yates and White, 2005) or when combined with the large ground reaction forces of running, increased hip joint compression. Repeated contractions with
Fig. 3. Amplitude of eversion. Mean and standard deviation of eversion with groups based on forefoot angle at contact (A) and based on rearfoot angle at contact (B).
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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Fig. 4. Mean trajectories (bold line) with standard deviations bands of groups indicating amplitude and duration of forefoot (top) and rearfoot (bottom) eversion. Vertical lines indicate the onset of inversion for the moderate (dashed) and large (solid) groups based on forefoot contact angle.
altered muscle torques present will result in altered loading at the hip (Horak, 2011) and these altered hip-loadings may contribute to the development of hip osteoarthritis and the need for hip replacement surgery (Gross et al., 2007; Horak et al., 2011). 5. Conclusion The results of this study indicate that an extrinsic non-weight bearing clinical measure of forefoot and rearfoot angles can predict, respectively, the forefoot and rearfoot angles at ground contact. However, the forefoot contact angle was associated with the amplitudes of both rearfoot and forefoot eversion during stance, whereas the rearfoot contact angle was associated with neither of these stance measures. These findings emphasize the importance of clinically obtaining an extrinsic measure of forefoot angle during nonweight bearing when 1) exploring the causes of non-traumatic lower extremity injuries in runners and 2) designing and prescribing foot orthoses to address foot structure. Acknowledgments Dudley Allen Sargent Grant, College of Health and Rehabilitation Sciences: Boston University. There was no involvement of sponsors in the study. References Brandt, K.D., Dieppe, P., Radin, E.L., 2008. Etiopathogenesis of osteoarthritis. Rheum. Dis. Clin. N. Am. 34 (3), 531–559. Buist, I., Bredeweg, S.W., Bessem, B., van Mechelen, W., Lemmink, K.A., Diercks, R.L., 2010. Incidence and risk factors of running-related injuries during preparation for a 4-mile recreational running event. Br. J. Sports Med. 44 (8), 598–604. Cole, G.K., Nigg, B.M., Ronsky, J.L., Yeadon, M.R., 1993. Application of the joint coordinate system to three-dimensional joint attitude and movement representation: a standardization proposal. J. Biomech. Eng. 115 (4A), 344–349. Davis, I., Ferber, R., Pollard, C.D., 2003. Rearfoot mechanics in competitive runners who had experienced plantar fasciitis. International Society of Biomechanics Conference. Dunedin, New Zealand. Ferber, R., Davis, I.M., Williams III, D.S., 2005. Effect of foot orthotics on rearfoot and tibia joint coupling patterns and variability. J. Biomech. 38 (3), 477–483.
Gross, K.G., Niu, J., Zhang, Y., Felson, D.T., Mc Lennan, C., Hannan, M.T., et al., 2007. Varus foot alignment and hip conditions in older adults. Arthritis Rheum. 6 (9), 2933–2938. Hamill, J., Bates, B., Knutzman, K., 1989. Relationship between selected static and dynamic lower extremity measures. Clin. Biomech. 4 (4), 417–425. Holt, K., Hamill, J. (Eds.), 1995. Running injuries and treatment: a dynamic approach. Sammaarco, G. (Ed.), 1995. Rehabilitation of the Foot and Ankle. Mosby, St Louis, pp. 241–257. Horak, Z., Kubovy, P., Stupka, M., Horakova, J., 2011. Biomechanical factors influencing the beginning and development of osteoarthritis in the hip joint. Wien. Med. Wochenschr. 161 (19–20), 486–492. Hsu, W.H., Lewis, C.L., Monaghan, G.M., Saltzman, E., Hamill, J., Holt, K.G., 2014. Orthoses posted in both forefoot and rearfoot reduce moments and angular impulses in lower extremity joints during walking. J. Biomech. (accepted for publication). Jacobs, S.J., Berson, B.L., 1986. Injuries to runners: a study of entrants to a 10,000 meter race. Am. J. Sports Med. 14 (2), 151–155. Korpelainen, R., Orava, S., Karpakka, J., Siira, P., Hulkko, A., 2001. Risk factors for recurrent stress fractures in athletes. Am. J. Sports Med. 29 (3), 304–310. Lafortune, M.A., Cavanagh, P.R., Sommer III, H.J., Kalenak, A., 1994. Foot inversion– eversion and knee kinematics during walking. J. Orthop. Res. 12 (3), 412–420. Lamppa, R., 2011. Annual half marathon reportAvailable http://www.runningusa.org/. Leardini, A., Benedetti, M.G., Berti, L., Bettinelli, D., Nativo, R., Giannini, S., 2007. Rear-foot, mid-foot and fore-foot motion during the stance phase of gait. Gait Posture 25 (3), 453–462. Lun, V., Meeuwisse, W.H., Stergiou, P., Stefanyshyn, D., 2004. Relation between running injury and static lower limb alignment in recreational runners. Br. J. Sports Med. 38 (5), 576–580. McClay, I., Manal, K., 1997. Coupling parameters in runners with normal and excessive pronation. J. Appl. Biomech. 13, 109–124. Michaud, T.C., 1993. Foot Orthoses and Other Forms of Conservative Foot Care. Williams & Wilkins, Baltimore, Maryland. Monaghan, G.M., Lewis, C.L., Hsu, W.H., Saltzman, E., Hamill, J., Holt, K.G., 2013. Forefoot angle determines duration and amplitude of pronation during walking. Gait Posture 38 (1), 8–13. Pohl, M., Buckley, J., 2008. Changes in foot and shank coupling due to alterations in foot strike pattern during running. Clin. Biomech. 23, 334–341. Razeghi, M., Batt, M., 2002. Foot type classification: a critical review of current methods. Gait Posture 15 (3), 282–291. Root, M., Orien, W., Weed, J. (Eds.), 1977. Normal & Abnormal Function of the Foot. Stackhouse, C., Davis Mc Clay, I., Hamill, J., 2004. Orthotic intervention in forefoot and rearfoot strike running patterns. Clin. Biomech. 19, 64–70. Surgeon General, 1999. Patterns and Trends in Physical Activity, Center for Chronic Disease Prevention and Health Wellness Promotion. http://www.cdc.gov/nccdphp/ sgr/contents.htm. Taunton, J.E., Ryan, M.B., Clement, D.B., McKenzie, D.C., Lloyd-Smith, D.R., Zumbo, B.D., 2002. A retrospective case–control analysis of 2002 running injuries. Br. J. Sports Med. 36 (2), 95–101. Tiberio, D., 1987. The effect of excessive subtalar joint pronation on patellofemoral mechanics: a theoretical model. J. Orthop. Sports Phys. Ther. 9 (4), 160–165.
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
G.M. Monaghan et al. / Clinical Biomechanics xxx (2014) xxx–xxx Van Gheluwe, B., Kirby, K.A., Roosen, P., Phillips, R.D., 2002. Reliability and accuracy of biomechanical measurements of the lower extremities. J. Am. Podiatr. Med. Assoc. 92 (6), 317–326. Williams III, D.S., McClay Davis, I., Baitch, S.P., 2003. Effect of inverted orthoses on lowerextremity mechanics in runners. Med. Sci. Sports Exerc. 35 (12), 2060–2068.
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Winter, D.A., 2005. Biomechanics and Motor Control of Human Movement, Third edition. John Wiley & Sons, Inc., University of Waterloo, Waterloo, Ontario, Canada. Yates, B., White, S., 2005. The incidence and risk factors in the development of medial tibial stress syndrome among naval recruits. Am. J. Sports Med. 33 (3), 463–464 (Mar).
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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Clinical Biomechanics journal homepage: www.elsevier.com/locate/clinbiomech
Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running Gail M. Monaghan a,⁎, Wen-Hao Hsu a, Cara L. Lewis a, Elliot Saltzman a, Joseph Hamill b, Kenneth G. Holt a a b
Department of Physical Therapy & Athletic Training, College of Health and Rehabilitation Sciences: Sargent College, Boston, MA, USA Department of Kinesiology, University of Massachusetts, Amherst, MA, USA.
a r t i c l e
i n f o
Article history: Received 25 September 2013 Accepted 19 June 2014 Keywords: Pronation Forefoot inversion Rearfoot inversion Musculoskeletal injury
a b s t r a c t Background: Clinically, foot structures are assessed intrinsically – relation of forefoot to rearfoot and rearfoot to leg. We have argued that, from a biomechanical perspective, the interaction of the foot with the ground may influence forces and torques that are propagated through the lower extremity. We proposed that a more appropriate measure is an extrinsic one that may predict the angle the foot makes with ground at contact. The purposes of this study were to determine if the proposed measure predicts contact angles of the forefoot and rearfoot and assess if the magnitude of those angles influences amplitude and duration of foot eversion during running. Methods: With the individual in prone, extrinsic clinical forefoot and rearfoot angles were measured relative to the caudal edge of the examination table. Participants ran over ground while frontal plane forefoot and rearfoot contact angles, forefoot and rearfoot eversion amplitude and duration were measured. Participants were grouped twice, once based on forefoot contact inversion angle (moderate b median and large N median) and once based on rearfoot contact inversion angle (moderate b median and large N median). Findings: The forefoot and rearfoot extrinsic clinical angles predicted, respectively, the forefoot and rearfoot angles at ground contact. Large forefoot contact angles were associated with greater amplitudes (but not durations) of forefoot and rearfoot eversion during stance. Rearfoot contact angles, however, were associated with neither amplitudes nor durations of forefoot and rearfoot eversion. Interpretation: Possible mechanisms for the increased risk of running injuries associated with large forefoot angles are discussed. © 2014 Elsevier Ltd. All rights reserved.
1. Introduction Thirty million Americans run to maintain health and fitness (Surgeon General, 1999). The typical runner has become more competitive, entering half-marathon races and logging in 1269 miles/year (Lamppa, 2011). One half of runners will experience musculoskeletal injuries which will stop them running for at least a week (Buist et al., 2010; Jacobs and Berson, 1986) and the cumulative effect of these injuries over many years of running may be knee and hip osteoarthritis (Brandt et al., 2008). As the foot contacts the ground many times, with ground reaction forces up to three times body weight, it is expected that foot structure will affect the stresses experienced not only by tissues reported as common running injury sites (i.e., patellofemoral joint, the iliotibial band and knee cartilage; (Taunton et al., 2002)), but also by the hip joint. Despite the prevalence of musculoskeletal injury
⁎ Corresponding author at: Boston University, College of Health and Rehabilitation Sciences: Sargent College, Department of Physical Therapy and Athletic Training, 635 Commonwealth Avenue, Boston, MA 02215, USA. E-mail address: [email protected] (G.M. Monaghan).
associated with running, the biomechanical mechanisms linking foot structure to abnormal movement patterns and potential injury are poorly understood (Razeghi and Batt, 2002). A critical factor that may contribute to this lack of understanding may be the way that foot structure is defined. Traditionally, foot structure is defined intrinsically, that is, relative to the other adjacent segments (forefoot to rearfoot, rearfoot to leg) (Root et al., 1977) (Fig. 1: Intrinsic measure). However, the intrinsic measure may not adequately represent the angle of the foot as it contacts the ground and, therefore, may not adequately capture the effects that contact angle may have on the forces and torques generated around the foot and upward through the kinematic chain. Additionally, given its poor inter-rater reliability (Van Gheluwe et al., 2002), using the intrinsic clinical measure of foot structure may contribute to the poor correlation of this measure to foot kinematics observed during the stance phase of locomotion (Hamill et al., 1989). Holt and Hamill (1995) proposed that the angle between the foot and the ground at foot contact during locomotion is the critical determinant of the force and torque patterns created, first at the foot and then through the rest of the body. In the current study of running, with the individual prone, clinical measures of rearfoot and forefoot angles
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Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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Fig. 1. Clinical foot measures. Intrinsic measure of clinical foot angle (left): rearfoot (B) and forefoot (A) angles are measured relative their respective proximal segments. The rearfoot angle is measured as the angle between the bisection of the calcaneus (black solid line) and the bisection of the lower leg (dashed black line). The forefoot angle is measured as the angle between the line through the metatarsal heads (black solid line) and the line bisecting the calcaneus. Extrinsic measure of clinical foot angle (right): rearfoot (C) and forefoot (D) angles are both measured relative to the caudal edge of the table. Rearfoot angle is measured as the angle between a line bisecting the calcaneus (white solid line) and a line perpendicular to the caudal edge of the table which is parallel to the longitudinal axis of the table (dashed black line). Forefoot angle is measured as the angle between the line connecting the first through fifth metatarsal heads (black solid line) and the caudal edge of the table (black dashed line).
were assessed extrinsically relative to a fixed environmental axis, the caudal edge of the table. This edge provides an environmental reference line that is functionally meaningful for locomotion in the sense that, were the table and subject to be rotated to an upright position, this edge would be parallel to the ground. Using this same method, Monaghan et al. (2013) previously showed that the angle of the forefoot to the ground at forefoot contact determines the amount and duration of eversion during walking, whereas the rearfoot angle at rearfoot contact had no effect on the amplitude or duration of eversion during walking. It is not known, however, if these findings also apply to running gait. Rearfoot kinematics during locomotion in the laboratory have been studied extensively (Davis et al., 2003; McClay and Manal, 1997; Pohl and Buckley, 2008) and have contributed to understanding injury mechanisms, but there is no epidemiological evidence to suggest that rearfoot angles measured clinically are associated with injury. Forefoot kinematics, however, have been less well studied. While some research has tested the application of foot orthoses to address clinically measured forefoot abnormalities (Ferber et al., 2005; Stackhouse et al., 2004; Williams et al., 2003), few studies report forefoot kinematics during locomotion or investigate the association between clinical forefoot abnormality and injury (Gross et al., 2007; Korpelainen et al., 2001; Lun et al., 2004). Holt and Hamill (1995) proposed that a clinical forefoot inversion deformity, measured in a non-weight bearing position, may be a prime contributor to the occurrence of pronation between heel off and toe off, which is referred to as “late” pronation. This late pronation could disrupt the coordinated movement of the foot, knee, and hip, and result in injury of proximal joints, a claim supported by others (Michaud, 1993; Tiberio, 1987). The effects of late forefoot pronation on rearfoot kinematics may be greater during running due to the presence of correspondingly larger ground reaction forces and resultant torques than during walking. Therefore, the purposes of this study were: 1) to test the ability of the extrinsic clinical measures of rearfoot and forefoot angles to predict the segment-to-ground angles of both rearfoot and forefoot at initial ground contact and 2) to explore the relationship between rearfoot and forefoot angles at initial contact and kinematic variables during the stance phase of running. 2. Methods 2.1. Subjects and instrumentation Eight females and six males (mean (SD) age 23.6 (4.9) years) were recruited for this study. The average and standard deviation mass and
height of participants were 66 (10.5) kg and 1.7 (0.1) m respectively. The inclusion criterion for all participants was to have an extrinsically measured forefoot angle greater than 15° of inversion. We selected this range because Gross et al. (2007) suggested that only the extremes of forefoot structural abnormality had an impact on injury. Individuals with current lower extremity pain were excluded. Prior to participation, each participant signed a consent form approved by the Institutional Review Boards of Boston University and University of Massachusetts, Amherst. Photographs of each subject's feet were taken using a six megapixel digital camera, Nikon Coolpix L11 which was fixed to the wall. Each subject was fitted with standardized sandals (Bite (LLC) running sandals with straps at the forefoot and rearfoot). An 8-camera Oqus motion capture system (Qualisys Medical AB, Gothenburg, Sweden) was used to record foot segment motions at 240 Hz. Ground reaction force data from a force plate (Advanced Mechanical Technologies, Inc., Watertown, MA, USA). Force plate data were synchronized with the motion data using Qualisys Track Manager. 2.2. Data collection On the first of three visits to the laboratory, the participant was screened to ensure that inclusion criteria were met. On the second visit, the participant received a clinical evaluation while lying in a prone position on an examination table. Briefly (see Monaghan et al., 2013), the medial malleoli were manually aligned at the caudal edge of the table, the legs were manually placed with the medial malleoli 33 cm apart and in neutral rotation, and, using light pressure at the third metatarsal head the examiner guided the ankle into 0° of dorsiflexion. The participant was instructed to hold the position after the examiner removed pressure from the third metatarsal head. A digital photograph was then taken. This positioning process was performed three times for each participant, and three digital pictures were obtained. Participants practiced running along a 10 m runway to ensure that the right foot contacted a force plate embedded in the runway. The participant's preferred speed was determined as the average speed from three such runs. In order to test the effect of small changes in running speed, the participant was then asked to run at slower and faster speeds than the preferred speed until consistent fast and slow speeds were reached. It was required that speed levels differ by at least 0.2 m/s. Only trials within approximately 0.1 m/s of the designated preferred, slow and fast speeds were accepted. Running speed was monitored using a set of photocell-triggered timers. Eleven retro-reflective
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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markers (9.5 mm in diameter) were placed on the foot and ankle according to the Leardini et al.'s model (2007), and a 12th ‘virtual marker’ was defined computationally as the midpoint between the sustentaculum tali and the peroneal tubercle markers (Fig. 2). Following an upright standing calibration trial, each participant performed five acceptable running trials at each of the three speeds. Approximately 1 month later on visit three, the motion capture procedures were repeated but without a clinical evaluation. 2.3. Data processing Rearfoot and forefoot extrinsic clinical angles were measured from the digital photographs using commercially available software (CanvasX, ACD Systems, Vancouver, British Columbia, Canada). The detailed protocol is described by Monaghan et al. (2013). Referencing the clinical angle to the environment (caudal edge of the table), rearfoot angle was defined as the angle between a line bisecting the calcaneus and a line perpendicular to the caudal edge of the table (Fig. 1C); forefoot angle was defined as the angle between the line connecting the first and fifth metatarsal heads and the caudal edge of the table (Fig. 1D). This edge defines a line that is parallel to the mediolateral axis in the participant's frontal
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plane and is parallel to the ground when standing. The average of measurements taken from the three photographs was used as the clinical extrinsic foot angle. The ICC's for intra-rater reliabilities were 0.91 for the forefoot and 0.87 for the rearfoot. Gross et al. (2007) reported similar intra-rater reliabilities, and also reported inter-rater reliabilities of 0.93 for the forefoot and 0.85 for the rearfoot. Marker data were labeled using Qualisys Track Manager. Raw kinematic data were low-pass filtered using a fourth order, zero-lag Butterworth digital filter with a cut-off frequency of 8 Hz. Three dimensional segment angles were calculated in Visual 3D using an x (flexion/extension), y (eversion/inversion), and z (axial rotation) Cardan rotation sequence (Cole et al., 1993). Eversion/inversion segment angles of the rearfoot and forefoot were calculated in the frontal plane relative to the horizontal and were used as proxies for pronation/supination. Using segment angles instead of joint angles assesses the foot motion relative to the ground. Time derivatives of the angular position time series were calculated in Visual 3D. Initial contact and toe-off were determined using the vertical ground reaction force. Stance phase data were normalized to percent stance duration. All runners were rearfoot strikers. Rearfoot contact was defined as initial contact; forefoot contact was defined when the third derivative (jerk) of the vertical position data of the fifth metatarsal head marker (VMH in Fig. 2) equaled zero, and heel off was assumed when the jerk of the vertical position data of the lateral heel marker (PT in Fig. 2) equaled zero (Winter, 2005). The onsets of rearfoot and forefoot inversion were defined when the velocities of the rearfoot and forefoot angles, respectively, in the frontal plane were greater than zero and remained above zero for the remainder of stance. Time series data for the segment angles were exported from Visual 3D. Custom Matlab (Mathworks, Natick, MA) programs extracted rearfoot and forefoot angle at foot contact, amplitude of rearfoot and forefoot eversion, duration of rearfoot and forefoot eversion, and duration (% stance) between heel off and the onset of rearfoot and forefoot inversion. Amplitude of rearfoot eversion was defined as the difference between the rearfoot angle at contact and its maximum eversion angle. Amplitude reflects the motion during the loading phase instead of focusing on a single time point. Duration of rearfoot eversion was defined as the percent of stance between contact and the onset of rearfoot inversion and therefore includes the time when the foot is neither inverting nor everting. Duration of forefoot eversion was defined similarly as the percent of stance between contact and the onset of forefoot inversion. Durations between heel off and the respective onsets of rearfoot and forefoot inversion were measured as percent of stance. 2.4. Participant categorization
Fig. 2. Foot markers, landmarks, and segments. In accordance with the model of Leardini et al. (2007) retro-reflective markers (solid circles) were placed on the bony landmarks of the foot: navicular (TN), first (FMB), second (SMB), and fifth (VMB) metatarsal bases, first (FMH), second (SMH) and fifth (VMH) metatarsal heads and the most distal and dorsal point of the head of the hallux (PM). Markers on the calcaneus were located at the triceps surae insertion (CA), the most medial apex of the sustentaculum tali (ST) and the lateral apex of the peroneal tubercle (PT). IC (gray circle) is intermedius calcaneus, an additional computed landmark (‘virtual marker’) defined at the midpoint between ST and PT. The rearfoot segment and forefoot segments are red dashed-line triangles. The rearfoot segment was defined by the triangle CA–ST–PT; its origin is located at CA, and its x-axis is defined along the line from the origin (CA) to IC. The forefoot segment was defined by the triangle SMB–FMH–VMH; its origin is located at SMB, and its x-axis is defined along the projection of the line from SMB to SMH onto the transverse plane passing through the origin (SMB), FMH and VMH. Image adapted from www.fotosearch.com. k8433317 “Bones of the Human Foot”. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
Participants in this study were grouped twice, once based on their forefoot contact angle and once based on the rearfoot contact angle. The groups were defined by dividing the participants into two equal groups based on the median contact angle: those participants above the median contact angle were in the “large inversion angle” group, and those below the median contact angle were in the “moderate inversion angle” group. This dual categorization allowed us to investigate the influence of contact angle on kinematic parameters separately for forefoot and rearfoot. We chose to compare moderate and large inversion angles, as only participants with very large forefoot inversion angles have been shown to be at risk of injury (Gross et al., 2007). This procedure allowed us to examine the relationship of forefoot and rearfoot contact angles with other foot-related kinematic measures obtained during the stance phase of running. 2.5. Statistical analysis A linear general estimating equation (GEE) analysis was used to determine if mean forefoot and rearfoot extrinsic clinical angles could predict mean forefoot and rearfoot angles at forefoot and rearfoot contact,
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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respectively, during running. Within-subject variables were speed (preferred, slow, fast) and visit (day 1 and day 2). To test if forefoot and rearfoot angles at contact predicted foot kinematics during stance, two separate three factor ANOVAs were performed – one for the groups based on forefoot angle at contact and one for the groups based on rearfoot angle at contact – in order to determine if there was a significant difference between groups in their predictions of foot kinematics during stance. The three factors were 1) group based on the magnitude of the contact angle (large and moderate), 2) running speed (preferred, slow, fast) and 3) visit (day 1 and day 2). The dependent variables were: 1) the amplitude of rearfoot eversion, 2) the amplitude of forefoot eversion, 3) the duration of rearfoot eversion (% stance), 4) the duration (% stance) of forefoot eversion, 5) the duration (% stance) between heel off and the onset of rearfoot inversion, and 6) the duration (% stance) between heel off and the onset of forefoot inversion. We were interested in the main effect of group and any interaction effect between group and speed or visit. Statistical significance was set at an alpha level of 0.05.
Table 1 Characteristics of groups based on contact angle (mean (SD)).
3. Results
moderate rearfoot inversion contact angle groups (Table 2, bottom) (Fig. 3B).
3.1. Effect of speed and visit There were no interaction effects between group and speed or group and visit. Therefore, we focused our analyses on the main effect of group. The average slow, preferred and fast speeds were 2.97, 3.48 and 4.19 m/s. 3.2. Relationships between clinical and contact angles 3.2.1. GEE analysis The P-value tests whether the clinical angle significantly predicts the contact angle, and the slope (B) can be used to describe how the expected (or mean) contact angle changes as the clinical angle score changes. The Wald statistic is a parametric statistic used to test how well the true value of a parameter can predict another parameter whenever the relationship between two parameters can be expressed as a model. The clinical forefoot angle predicted the forefoot angle at contact during running with a positive linear relation (B = 0.14, W(1, 11) = 5.1, P = 0.02). The clinical rearfoot angle also predicted the rearfoot angle at contact, but with a negative linear relation indicating an inverse relationship (B = −.18, W(1, 7) = 25.4, P ≤ 0.01). There was no effect of running speed for the forefoot contact angle. Rearfoot contact angle increased significantly as the speed increased (mean 3.54° [SD 0.48°], mean 4.55° [SD 0.43°], mean 4.60° [SD 0.51°], P = 0.01). 3.2.2. Group comparisons The group with large forefoot inversion angles at forefoot contact had larger clinical angles (Table 1) than the group with moderate forefoot inversion angles (F(1,12) = 11.7, P = 0.01), reflecting the positive linear relation between forefoot clinical angle and forefoot contact angle reported in the regression analysis. Conversely, the group with large rearfoot inversion contact angles at rearfoot contact had smaller clinical angles than the group with moderate rearfoot inversion contact angles (F(1,8) = 6.7, P = 0.03), reflecting the negative linear relation between rearfoot clinical angle and rearfoot contact angle reported in the regression analysis. 3.3. Relationship of contact angles to foot motion 3.3.1. Amplitude of eversion The large forefoot inversion contact angle group had greater amplitude of forefoot and rearfoot eversion (Table 2) (Fig. 3A) than the moderate forefoot inversion group. There were no differences in amplitudes of forefoot or rearfoot eversions between the large rearfoot and the
Forefoot grouping
Moderate
Large
Forefoot angle at contact Rearfoot angle at contact Age Weight Height Forefoot clinical angle Rearfoot clinical angle
3.2° (1.7°) 3.8° (2.4°) 24.9 (6.5) years 70.1 (13.4) kg 1.7 (0.12) m 19.4° (7.1°) 3.6° (10.1°)
7.0° (1.0°) 5.2° (2.0°) 22.4 (2.7) years 63.1 (9.9) kg 1.7 (0.14) m 28° (2.7°) 5.7° (5.2°)
Rearfoot grouping
Moderate
Large
Rearfoot angle at contact Forefoot angle at contact Age Weight Height Forefoot clinical angle Rearfoot clinical angle
2.9° (1.4°) 4.9° (3.6°) 21 (1.4) years 63.3 (15.1) kg 1.6 (0.12) m 27° (4.3°) 9.9° (7.1°)
6.1° (1.4°) 6.1° (3.1°) 22.4 (2.4) years 68.2 (12.7) kg 1.7 (0.13) m 20.6° (8.3°) −0.9° (6.2°)
3.3.2. Duration of eversion There were no differences between groups based on either forefoot or rearfoot contact angle in the duration of forefoot or rearfoot eversion (Table 2). 3.3.3. Duration of stance between heel off and onset of inversion There were no differences between groups based on either forefoot or rearfoot contact angle in the durations of stance between heel off and the onsets of both forefoot inversion and rearfoot inversion (Table 2). 4. Discussion The first goal of this study was to examine the predictive validity of a hypothesized extrinsically defined clinical foot angle measured during non-weight bearing. It was shown that forefoot angle, relative to a fixed environmental reference plane in non-weight bearing, can predict frontal plane angle of the forefoot at contact when running. This new extrinsic measure of foot angle showed predictive validity and had high intra-rater reliability for the forefoot and the rearfoot. However, the finding that a greater clinical rearfoot inversion angle measured in non-weight bearing predicted a smaller rearfoot angle at contact calls into question the common clinical assumption that the non-weight bearing measure of rearfoot is a critical factor in causing excessive pronation and for designing orthoses. According to Holt and Hamill (1995) and the findings of Lafortune et al. (1994), a large forefoot inversion angle at ground contact can produce large pronation torques that result in increased forefoot and rearfoot eversion. The finding in the current study of greater amplitude of forefoot and rearfoot eversion for individuals with a large clinical forefoot angle supports these claims (Fig. 4). Despite the emphasis on the rearfoot in orthotic prescription and in musculoskeletal injury research, rearfoot angle at contact did not appear to be associated with any of the kinematic variables measured during stance. The amplitude and duration of rearfoot eversion may have been influenced by forefoot amplitude and thus masked the association between the rearfoot contact angles and rearfoot kinematic variables. In this paper, we have shown that the foot orientation at ground contact has an important influence on foot biomechanics. This is a unique finding because it provides evidence to support the hypothesis that foot structure and its biomechanical interaction with the ground is a critical first step when investigating the causes of foot and proximal injuries. To fully support the relationship between foot structure and injury, a second step is to show that the high amplitude and velocity of
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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Table 2 Differences in foot contact groups: mean, SD, F ratios and P values for each dependent variable. Variable
Amplitude of forefoot eversion (°) Amplitude of rearfoot eversion (°) Duration of forefoot eversion (%) Duration of rearfoot eversion (%) Duration between heel off and onset of forefoot inversion (%) Duration between heel off and onset of rearfoot inversion (%)
Group
Effect size
Moderate
Large
FF group
Mean (SD)
Mean (SD)
F(11)
P
ES
7.6 (1.1) 3.7 (0.4) 44.9 (1.5) 52.1 (6.1) 13.9 (2.0) 24.5 (4.0)
11.6 (1.1) 6.1 (0.4) 44.5 (1.5) 65.2 (5.5) 12.1 (2.1) 32.6 (3.6)
7 20 0.04 2.6 0.42 5.6
0.02* b0.01* 0.85 0.15 0.53 0.16
3.6 6 −0.26 0.37 −0.88 2.1
Moderate
Large
RF group effect
Mean (SD)
Mean (SD)
F(8)
10.3 (1.7) 5.0 (0.7) 44.8 (1.6) 58.7 (6.4) 14.0 (2.1) 29.6 (4.1)
10.1 (1.7) 4.8 (0.7) 43.1 (1.7) 60.2 (7.1) 11.0 (2.1) 28.0 (4.6)
Differences in rearfoot contact groups: mean, SD, F ratios and P values for each dependent variable Variable
Amplitude of forefoot eversion (°) Amplitude of rearfoot eversion (°) Duration of forefoot eversion (%) Duration of rearfoot eversion (%) Duration between heel off and onset of forefoot inversion (%) Duration between heel off and onset of rearfoot inversion (%)
Group
segment eversion found in this study are associated with changes in proximal motions and torques. We have previously shown that a forefoot and rearfoot posted foot orthotic reduces the motions and torques at the foot and in the ankle, knee and hip during walking (Hsu et al., accepted for publication, J. Biomechanics). A final step is to show how these stresses result in tissue injuries. While our epidemiological study has shown a 5-times greater occurrence of hip replacements in elderly individuals with large forefoot varus (Gross et al., 2007), a similar study is needed for a large group of runners. In addition to these studies, it will also be important to investigate the interaction between foot orientation, and lower extremity joint angles (including forefoot to rearfoot) at foot contact. For example, in addition to foot orientation, the influence of frontal plane shank position relative to the foot will likely influence the motion and torques at the ankle and leg. When taken together, the foot contact angle and joint angles will present a more complete picture of the causes of non-traumatic lower extremity musculoskeletal injury. The comparison of groups with moderate and large amounts of inversion abnormality provided the most stringent of tests for the effects of structure on the kinematics of the foot during running. Nevertheless, it would be interesting and informative to have included a group with minimal or no inversion deformity to compare our groups to ‘normal’ feet. Furthermore, in this study all subjects had a prior history of some
Effect size
0.01 0.05 0.71 0.03 1.3 0.07
P
ES 0.94 0.83 0.42 0.88 0.28 0.79
−0.12 −0.29 −1 0.22 −1.42 −0.37
form of injury either to the feet or lower extremity. However, we made no attempt to investigate the differential effects of type of foot structure associated with specific injuries, particularly to more proximal structures. In future studies we plan to a) investigate the effects of foot structure on the biomechanics of more proximal segments, and b) explore the reasons that different forms of chronic, non-traumatic injury may be related to similar structural abnormalities. The finding that individuals with a greater forefoot inversion angle at contact also had greater forefoot and rearfoot eversion amplitude but not greater eversion duration suggests a mechanism for the occurrence of stresses in distal or proximal regions of the lower extremity. If greater pronation amplitude occurs in the same duration of pronation as individuals with a lesser forefoot inversion angle at contact, then individuals with a greater inversion contact angle had greater forefoot and rearfoot eversion velocities. Subsequently there will be greater tibial and femoral internal rotation velocity and amplitude. The foot of a runner contacts the ground many times each year. For a runner with a large forefoot inversion angle muscles which slow rearfoot eversion and femoral and tibial rotation will have to work harder to control faster movement of greater amplitude. The greater torque generated by these muscles may result in soft tissue injuries (Yates and White, 2005) or when combined with the large ground reaction forces of running, increased hip joint compression. Repeated contractions with
Fig. 3. Amplitude of eversion. Mean and standard deviation of eversion with groups based on forefoot angle at contact (A) and based on rearfoot angle at contact (B).
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
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G.M. Monaghan et al. / Clinical Biomechanics xxx (2014) xxx–xxx
Fig. 4. Mean trajectories (bold line) with standard deviations bands of groups indicating amplitude and duration of forefoot (top) and rearfoot (bottom) eversion. Vertical lines indicate the onset of inversion for the moderate (dashed) and large (solid) groups based on forefoot contact angle.
altered muscle torques present will result in altered loading at the hip (Horak, 2011) and these altered hip-loadings may contribute to the development of hip osteoarthritis and the need for hip replacement surgery (Gross et al., 2007; Horak et al., 2011). 5. Conclusion The results of this study indicate that an extrinsic non-weight bearing clinical measure of forefoot and rearfoot angles can predict, respectively, the forefoot and rearfoot angles at ground contact. However, the forefoot contact angle was associated with the amplitudes of both rearfoot and forefoot eversion during stance, whereas the rearfoot contact angle was associated with neither of these stance measures. These findings emphasize the importance of clinically obtaining an extrinsic measure of forefoot angle during nonweight bearing when 1) exploring the causes of non-traumatic lower extremity injuries in runners and 2) designing and prescribing foot orthoses to address foot structure. Acknowledgments Dudley Allen Sargent Grant, College of Health and Rehabilitation Sciences: Boston University. There was no involvement of sponsors in the study. References Brandt, K.D., Dieppe, P., Radin, E.L., 2008. Etiopathogenesis of osteoarthritis. Rheum. Dis. Clin. N. Am. 34 (3), 531–559. Buist, I., Bredeweg, S.W., Bessem, B., van Mechelen, W., Lemmink, K.A., Diercks, R.L., 2010. Incidence and risk factors of running-related injuries during preparation for a 4-mile recreational running event. Br. J. Sports Med. 44 (8), 598–604. Cole, G.K., Nigg, B.M., Ronsky, J.L., Yeadon, M.R., 1993. Application of the joint coordinate system to three-dimensional joint attitude and movement representation: a standardization proposal. J. Biomech. Eng. 115 (4A), 344–349. Davis, I., Ferber, R., Pollard, C.D., 2003. Rearfoot mechanics in competitive runners who had experienced plantar fasciitis. International Society of Biomechanics Conference. Dunedin, New Zealand. Ferber, R., Davis, I.M., Williams III, D.S., 2005. Effect of foot orthotics on rearfoot and tibia joint coupling patterns and variability. J. Biomech. 38 (3), 477–483.
Gross, K.G., Niu, J., Zhang, Y., Felson, D.T., Mc Lennan, C., Hannan, M.T., et al., 2007. Varus foot alignment and hip conditions in older adults. Arthritis Rheum. 6 (9), 2933–2938. Hamill, J., Bates, B., Knutzman, K., 1989. Relationship between selected static and dynamic lower extremity measures. Clin. Biomech. 4 (4), 417–425. Holt, K., Hamill, J. (Eds.), 1995. Running injuries and treatment: a dynamic approach. Sammaarco, G. (Ed.), 1995. Rehabilitation of the Foot and Ankle. Mosby, St Louis, pp. 241–257. Horak, Z., Kubovy, P., Stupka, M., Horakova, J., 2011. Biomechanical factors influencing the beginning and development of osteoarthritis in the hip joint. Wien. Med. Wochenschr. 161 (19–20), 486–492. Hsu, W.H., Lewis, C.L., Monaghan, G.M., Saltzman, E., Hamill, J., Holt, K.G., 2014. Orthoses posted in both forefoot and rearfoot reduce moments and angular impulses in lower extremity joints during walking. J. Biomech. (accepted for publication). Jacobs, S.J., Berson, B.L., 1986. Injuries to runners: a study of entrants to a 10,000 meter race. Am. J. Sports Med. 14 (2), 151–155. Korpelainen, R., Orava, S., Karpakka, J., Siira, P., Hulkko, A., 2001. Risk factors for recurrent stress fractures in athletes. Am. J. Sports Med. 29 (3), 304–310. Lafortune, M.A., Cavanagh, P.R., Sommer III, H.J., Kalenak, A., 1994. Foot inversion– eversion and knee kinematics during walking. J. Orthop. Res. 12 (3), 412–420. Lamppa, R., 2011. Annual half marathon reportAvailable http://www.runningusa.org/. Leardini, A., Benedetti, M.G., Berti, L., Bettinelli, D., Nativo, R., Giannini, S., 2007. Rear-foot, mid-foot and fore-foot motion during the stance phase of gait. Gait Posture 25 (3), 453–462. Lun, V., Meeuwisse, W.H., Stergiou, P., Stefanyshyn, D., 2004. Relation between running injury and static lower limb alignment in recreational runners. Br. J. Sports Med. 38 (5), 576–580. McClay, I., Manal, K., 1997. Coupling parameters in runners with normal and excessive pronation. J. Appl. Biomech. 13, 109–124. Michaud, T.C., 1993. Foot Orthoses and Other Forms of Conservative Foot Care. Williams & Wilkins, Baltimore, Maryland. Monaghan, G.M., Lewis, C.L., Hsu, W.H., Saltzman, E., Hamill, J., Holt, K.G., 2013. Forefoot angle determines duration and amplitude of pronation during walking. Gait Posture 38 (1), 8–13. Pohl, M., Buckley, J., 2008. Changes in foot and shank coupling due to alterations in foot strike pattern during running. Clin. Biomech. 23, 334–341. Razeghi, M., Batt, M., 2002. Foot type classification: a critical review of current methods. Gait Posture 15 (3), 282–291. Root, M., Orien, W., Weed, J. (Eds.), 1977. Normal & Abnormal Function of the Foot. Stackhouse, C., Davis Mc Clay, I., Hamill, J., 2004. Orthotic intervention in forefoot and rearfoot strike running patterns. Clin. Biomech. 19, 64–70. Surgeon General, 1999. Patterns and Trends in Physical Activity, Center for Chronic Disease Prevention and Health Wellness Promotion. http://www.cdc.gov/nccdphp/ sgr/contents.htm. Taunton, J.E., Ryan, M.B., Clement, D.B., McKenzie, D.C., Lloyd-Smith, D.R., Zumbo, B.D., 2002. A retrospective case–control analysis of 2002 running injuries. Br. J. Sports Med. 36 (2), 95–101. Tiberio, D., 1987. The effect of excessive subtalar joint pronation on patellofemoral mechanics: a theoretical model. J. Orthop. Sports Phys. Ther. 9 (4), 160–165.
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
G.M. Monaghan et al. / Clinical Biomechanics xxx (2014) xxx–xxx Van Gheluwe, B., Kirby, K.A., Roosen, P., Phillips, R.D., 2002. Reliability and accuracy of biomechanical measurements of the lower extremities. J. Am. Podiatr. Med. Assoc. 92 (6), 317–326. Williams III, D.S., McClay Davis, I., Baitch, S.P., 2003. Effect of inverted orthoses on lowerextremity mechanics in runners. Med. Sci. Sports Exerc. 35 (12), 2060–2068.
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Winter, D.A., 2005. Biomechanics and Motor Control of Human Movement, Third edition. John Wiley & Sons, Inc., University of Waterloo, Waterloo, Ontario, Canada. Yates, B., White, S., 2005. The incidence and risk factors in the development of medial tibial stress syndrome among naval recruits. Am. J. Sports Med. 33 (3), 463–464 (Mar).
Please cite this article as: Monaghan, G.M., et al., Forefoot angle at initial contact determines the amplitude of forefoot and rearfoot eversion during running, Clin. Biomech. (2014), http://dx.doi.org/10.1016/j.clinbiomech.2014.06.011
