The heel pad is a highly specialized fibroadipose tissue that is believed to play an important protective role during gait, by minimizing local peak stresses during weight-bearing and dissipating transients generated during heel strike. Consisting of organized fibrous compartments that envelope and retain adipose tissue, the heel pad can be anatomically divided into a superficial microchamber layer and a deep macrochamber layer (15). Although both layers contribute to the properties of the tissue as a whole, the bulk mechanical properties of the heel pad predominantly reflect those of the collagen- and elastin-rich septal walls of the macrochamber layer, which envelope adipocytes and confine their movement (13). The altered mechanical properties of the heel pad have been reported with aging and several disease states (16,29,39). However, there is little consensus regarding the structural properties of the heel pad, especially under loading conditions encountered during activities of everyday living, such as walking.
To date, the structural properties of the human heel pad have been evaluated by either mechanical testing of cadaveric heels in vitro (1–3,5) or impact testing of heels in vivo (1,17,42) and have mostly simulated loading conditions associated with running. The reported stiffness of cadaveric heel pads under such conditions and at loads equivalent to bodyweight (700–800 N) range between 905 and 1445 N·mm−1, with 28%–48% of the energy required to compress the heel pad dissipated during each cycle of loading and unloading (2,3,5). In vivo measurements, during impact testing, in contrast, have typically yielded heel pad stiffness values 10 times lower (50–150 N·mm−1) with substantially higher energy losses (72%–99%) (1,7,17,37,42). Although such discrepancies in heel pad properties have been attributed to artifacts associated with impact testing and the damping properties of other structures of the leg (2,3,42), Bennett and Ker (5) reported average peak deformations of only 2.0 ± 0.3 mm with mechanical testing of cadaveric heel pads to loads equivalent to bodyweight. Such deformations with mechanical testing are approximately fivefold less than that observed in one of the few studies that have evaluated the in vivo deformation of the heel during gait (8). Hence, mechanical testing protocols do not seem to accurately replicate in vivo loading of the heel during gait.
In a study using cine-fluoroscopy to evaluate the in vivo deformation of the human heel pad during running, De Clercq et al. (8) argued that the heel pad was maximally deformed (≈60%) during barefoot running (4.5 m·s−1), in which impact transients between 1.9 and 2.9 times bodyweight were recorded. As such, the authors concluded that rather than attenuate peak forces associated with heel contact, the heel pad primarily acted as local protection to the tuber calcaneum. The suggestion that the heel pad operated at its “physiological maximum” during running is consistent with later observations that individuals make active gait adjustments that result in a more plantar-grade foot placement during barefoot running, thereby effectively lowering the peak pressure beneath the heel pad (9,21).
Although it has been suggested that the heel pad operates below this physiological threshold at slower gait speeds (11,44), this has not been empirically tested during walking. Slower walking speeds lead to lower peak pressures beneath the heel (6) and would intuitively result in less compression of the heel pad. However, the deformation behavior of the heel pad is nonlinear in vivo and loading rate dependent (1). Even quasistatic loading protocols, in which the heel pad was slowly loaded to bodyweight during unipedal stance, have noted similar absolute deformations of the heel pad (10–11 mm) to that reported during running (31,35), despite peak forces that were two to three times lower (8). Therefore, it is possible that the heel pad is physiologically optimized to operate at compressive strains of approximately 60% even at slower gait speeds, such as those encountered during walking. To date, however, the deformation and the structural properties of the heel pad during walking have received little empirical attention. The aim of the current investigation, therefore, was to evaluate and describe the deformation properties of the heel pad in healthy human adults while walking at their preferred gait speed. We hypothesized that a near-maximal deformation of the heel pad would also occur during walking, as has been shown previously with running (8).
A convenience sample of 10 healthy adults (6 males and 10 females, age = 45 ± 10 yr; height = 1.66 ± 0.10 m, and weight = 80.7 ± 10.8 kg; mean ± SD) free of gross orthopedic deformity of the lower extremities and feet participated in the study. No participant reported a medical history of endocrine disorders, inflammatory joint disease, lower limb surgery, plantar heel pain, or trauma of the foot. Participants gave their written informed consent to the procedures of the study, which received approval from the institutional ethical committee review board.
Dynamic lateral foot radiographs and plantar pressure data were collected as participants walked barefoot at their preferred speed using previously described methods (39). In brief, data were collected for three walking trials using a midgait protocol, which included a preamble of at least three steps and was conducted at a self-selected speed (41). Consistency between walking trials was ensured by monitoring the stance phase duration of each footstep, which differed by less than 5%. Trials were omitted if footsteps did not fall entirely within the boundaries of the fluoroscopic field of view, or if the investigators observed gait adjustments secondary to visual targeting of the platform. No participant reported pain or discomfort during the barefoot walking trials.
Foot radiographs were acquired using a C-Vision multifunction fluoroscopy unit, configured with a 40.5-cm four-field image intensifier (Shimadzu, Kyoto, Japan). Dynamic images (1024 × 1024 pixels) were acquired at a rate of 15 Hz, with an intensity of 50 kV and a radiation exposure equivalent to 1.2 mA·s−1. Spatial distortion within the imaging system was minimized using a rectilinear calibration grid (32 × 32 cm) positioned within the object plane and perpendicular to the central ray of the fluoroscope, in combination with a distortion correction procedure (38). The following application of convolution and edge detection algorithms to enhance the bone–soft tissue interface (Matlab software; MathWorks Inc., Natick, Massachusetts) was applied to manually digitize the inferior aspect of the calcaneus from initial heel contact until heel lift. The unloaded sagittal thickness and deformation of the heel pad relative to the support surface were subsequently calculated (Fig. 1). The root-mean-square error for repeated linear measures using this method is less than 0.1 mm (38).
An EMED-SF capacitance mat transducer system (Novel GmbH, Munich, Germany), with an effective spatial resolution of four sensors per square centimeters and mounted within the field of view, was used to simultaneously acquire barefoot plantar pressure data. The force beneath the heel was estimated using a standardized masking procedure, in which the length of the maximum pressure footprint, excluding the toes, was divided into equal thirds. The instantaneous force beneath the heel was calculated as the product of the area of active pressure sensors and the mean pressure within the rear third of the footprint. Force data were resampled from 50 to 15 Hz, and force–deformation data for each gait trial were subsequently plotted and principal structural properties, including the peak force and the peak deformation of the heel pad, were calculated. The force–deformation curve involved both loading and unloading curves. The loading curve reflected the deformation of the heel pad from initial contact to the time of peak force beneath the heel, whereas the unloading curve reflected the restitution of the heel pad from the time of peak force until heel lift. In accordance with previous research (7), the initial stiffness of the heel pad, corresponding to the slope of the force–deformation curve at forces less than 250 N, was estimated from loading curve data. Similarly, the final stiffness of the heel pad was calculated as the slope of the force–deformation curve at forces higher than 250 N (7). In each instance, an iterative linear regression approach, in which the least square error was minimized, was used to calculate stiffness more than 40% of the corresponding segment of the loading curve.
The area beneath the loading curve, reflecting the work done in compressing the heel pad, was calculated by numerical integration, as was the area under the unloading curve. The difference in area between the loading and unloading curves, representing the area of the hysteretic loop, was used to calculate the energy dissipation ratio (EDR) of the heel pad. The EDR was defined as the area of the hysteretic loop relative to the area under the loading curve.
In combination with previously published data (1,7,8,16,17,36,42), the deformation of the heel pad was subsequently plotted as a function of the work required to deform the fat pad. The viscoelastic properties of human soft tissues have been extensively modeled using both power-law and single-term exponential functions. Although exponential functions are prominent within the literature, the power law has a long history in tissue mechanics and has been successfully used to model the viscoelastic properties of the human calcaneal fat pad (10). Hence, combined data were modeled by a power function in which the deformation of the heel pad (D) was given by D(w) = awb, where w represents the energy to deform the heel pad and scaling factors a and b (representing heel pad thickness at 1 J and rate of deformation, respectively) were estimated using nonlinear regression and minimizing root mean square error. The deformation of the heel pad at an energy of 2.12 J, the upper limit of pain tolerance reported during impact testing of human heel pads in vivo (7), was subsequently estimated from the power function.
The Statistical Package for the Social Sciences (SPSS Inc., Chicago, IL) was used for all statistical procedures. The Kolmogorov–Smirnov tests were used to evaluate data for underlying assumptions of normality. Outcome variables were determined to be normally distributed, so mean and SD values were used as summary statistics. Potential differences between initial and final stiffness of the heel pad were investigated using a paired t-test, in which an alpha level of 0.05 was used for tests of significance.
Transient loading profiles associated with walking induced rapidly changing deformation rates in the heel fat pad and created irregular force–deformation curves (Fig. 2). With increasing load, the deformation of the fibroadipose tissue was relatively linear at loads less than 250 N. Above this load, the rate of heel pad deformation quickly reduced, and the vertical stiffness of the heel pad increased markedly. This characteristic point of inflection observed in the loading phase of the force–deformation curve occurred between 220 and 350 N in all cases.
The major structural properties of the heel fat pad while walking at preferred speed are summarized in Table 1. The average stance phase duration during walking was 930 ± 110 ms. On average, the heel pad was exposed to a force of 52% ± 4% bodyweight during walking and was compressed by 10.3 ± 1.9 mm from an unloaded thickness of 18.9 ± 1.7 mm to a final loaded thickness of 8.6 ± 1.6 mm (t15 = 19.1, P < 0.05). The initial stiffness of the heel pad during loading was 10 times lower than its final stiffness (t15 = −5.9, P < 0.05). The mean EDR of the heel pad during walking was 0.66 ± 0.12, with approximately 1 ± 0. 2 J of energy dissipated by the heel pad with each step.
Figure 3 shows the average deformation of the heel pad as a function of the work required to deform the heel fat pad in the current study, compared with that of other in vivo studies reported within the literature. When taken collectively, the relationship between input energy and heel pad deformation was best defined by a power function (coefficient of determination = 0.652), in which heel pad deformation approaches a plateau of 12 mm at an input energy of around 5.0 J. The mean work to deform the heel pad during walking in the current study was 1.52 ± 0.27 J. The limit of pain tolerance reported for in vivo impacts of the heel pad (2.12 J) (7) corresponds to a predicted heel pad deformation of 10.7 mm, which is marginally greater than the mean deformation value noted in the current study (10.3 ± 1.9 mm) for subjects walking at their preferred speed.
This study evaluated the deformation properties of the heel pad during walking in healthy human adults as opposed to those reported during running and impacts simulating running. On the basis of the work of Nilsson et al. (27), the average stance phase duration for the participants in this study, equated to a walking velocity of 0.9–1.0 m·s−1, which is consistent with a “slow” walking speed reported for similarly aged individuals (28). The transient loading profiles associated with this slow walking induced rapidly changing deformation rates in the heel fat pad and created irregular force–deformation curves characterized by a point of inflection of the loading phase curve at around 250 N. This finding is consistent with the point of inflection (or so-called knee) observed by Cavanagh et al. (7) during impact testing of the heel pad in vivo (1.61 J). However, the initial (32 N·mm−1) and the final (211 N·mm−1) stiffness of the heel pad in the current study were approximately twice those reported by Cavanagh et al. (7) during impact testing (19 and 138 N·mm−1, respectively). It is notable that the reported elastic (Young’s) modulus of the deep macrochambers of the heel pad (46 ± 18 kPa) has also been estimated to be 10 times lower than the adjacent superficial microchamber layer (450 ± 240 kPa) during indentation testing in vivo (13). Although further research is required, the initial and final stiffness of the overall force–deformation curve may reflect the structural properties of each of the component layers of heel pad; suggesting the heel pad may be viewed as a series elastic structure.
As with most soft biological tissues, a proportion of the strain energy stored within the heel fat pad is dissipated during gait. In the current study, the average EDR of the heel pad was 0.66, indicating that approximately 1.0 J of energy was dissipated by the fat pad during each step. This corresponds to approximately 1% of the total energy exchanged during a single gait cycle (~100 J) (4). Although EDR in the order of 0.39–0.67 are typically reported after a single cycle of loading and unloading of cadaveric heel pads (2,3,29), losses between 0.50 and 0.70 (39) and as low as 0.17 (11) have been reported in previous radiokinematic studies investigating the loss characteristics of the heel pad during walking. Although reasons underlying the marked departure in energy dissipation values noted in these latter studies are unclear, one possibility is that it may reflect the roughly fourfold difference in peak stress reported beneath the heel (≈70 vs ≈250 kPa). In the current study, the peak force beneath the heel (450 N) was comparable with that reported previously in healthy young adults walking at their preferred gait speed (500 N) (41) but substantially lower than that typically used during materials testing of cadaveric heel pads in vitro (1000–2000 N) (2,5). Thus, a direct comparison with these latter materials testing studies is not recommended and may not be physiologically relevant for activities such as walking and slow running (3.3 m·s−1), in which loads beneath the heel do not typically exceed 80% bodyweight (22,40). Interestingly, the EDR of the heel pad during walking in the current study is comparable with that reported during closely controlled impact testing of cadaveric heel pads (0.67), when performed using similar loading conditions (impact energy of 1.45 J and peak force of 450 N ), highlighting the importance of loading conditions in evaluating the mechanical properties of the heel pad.
As illustrated in Figure 3, the relationship between contact energy and heel pad deformation is best defined by a power function, in which heel pad deformation approaches a plateau of 12 mm at an input energy of approximately 5.0 J. Although indicative of an upper mechanical limit, beyond which the heel pad is no longer deformable (i.e., “bottoms out”), such a limit is unlikely to be realized during normal physiological loading. The limit of pain tolerance for impacts involving the heel pad occurs at energy levels higher than 2.12 J (7) and corresponds to a predicted heel pad deformation of 10.7 mm, which is marginally greater than mean deformation values noted in the current study (10.3 ± 1.9 mm). This finding suggests that, even at preferred walking speeds, the deformation of the heel fat pad during barefoot walking approaches the limits of pain tolerance. Such a finding has important clinical implications when walking barefoot, particularly at speeds greater than self-selected walking speed. For instance, to avoid potential pain and injury at higher locomotor speeds, either the contact energy must remain unchanged or the movement of the rearfoot and soft tissues of the shank must also contribute to energy dissipation. There is some evidence to suggest that soft tissues of the leg may enhance dissipation by exerting a “wobbling mass effect” (30). However, the energy required to deform the heel pad during walking in the current study (1.5 ± 0.3 J) was only marginally less than that reported during barefoot running (1.8 J) at moderate speeds (4.5 m·s−1) (8), suggesting that individuals maintain a similar contact energy at higher gait speeds. Hence, active kinematic adjustments that reduce the effective mass of the leg and foot likely take place at speeds faster than walking, which is consistent with the so-called Robbins and Hanna hypothesis, in which plantar foot sensation was proposed to moderate impact loading behavior during gait (32). Consequently, the observation that the heel pad operates close to its pain threshold during slow barefoot walking may also, in part, account for the reported “forefoot” strike pattern adopted during barefoot running at higher gait speeds (21) and reflect a pain-avoidance strategy.
The mechanical advantage of deforming the heel pad to a constant limit, even at the modest gait speeds encountered in this study, is not well understood. Speculatively, the near-constant deformation of the heel pad over a range of physiologically relevant speeds may incur a somatosensory benefit and play an important role in the regulation of gait. The heel pad is known to be richly innervated with Pacinian corpuscles, low-threshold encapsulated receptors that are localized to the deeper components of the fat pad (15). Such receptors are known to be sensitive to vibrations in the range of 60–300 Hz (34), and thresholds mediated by the Pacinian channel are dependent on both the strain rate and the final compressive strain of the tissue (12,18). Pacinian thresholds have also been shown to be lowered with larger areas of stimulation by spatial summation (24,25). Thus, the deformation of the heel pad to a constant strain over a range of speeds would ensure a consistent contact area of the heel pad and potentially maximize the sensitivity of the heel pad to vibration associated with heel strike. Such a mechanism is consistent with observations that plantar pressures are increased beneath the heel with faster walking speeds while the contact area remains unaltered (6). Such a mechanism may also underpin the active kinematic adjustments observed during shod running with variations in midsole hardness, which act to maintain consistent external vertical impact forces (26). As such, research evaluating the potential neuromechanical role of the heel fat pad seems warranted.
A limitation of the current study is that the temporal resolution of the imaging system resulted in a sampling rate of 15 Hz, which is lower than that of modern motion analysis systems. Although impulsive transients as high as 100 Hz have been recorded during heel strike (33), the sampling rate in the current study is sufficient to capture the majority of skeletal movement during gait, in which 99% of the power of kinematic signals occurs at frequencies less than 6 Hz (43). An additional limitation of fluoroscopic imaging is that it produces relatively low contrast between different soft tissue structures, and as such, the contribution of various internal structures, such as the superficial or deep macrochambers, to the behavior of the heel pad cannot be independently determined. Similarly, fluoroscopy is a transmission technique that creates a two-dimensional projection of what is undoubtedly a three-dimensional deformation, and the pressure platform used in this study provided an estimate of only the vertical ground reaction force beneath the heel. Consequently, the force–deformation properties of the heel pad were idealized to only one (compression) dimension. There is evidence, however, that the properties of the heel pad are likely isotropic in nature (23) and that the magnitude of tangential force components at the heel are <10% of the peak vertical force during barefoot walking (20).
The present experimental approach did not allow for the quantification of the duration of double support or unloading of the contralateral limb during the time of heel contact. We would anticipate that the asymmetric loading during the initial period of double support would influence the rate and magnitude of loading of the heel pad. However, the period of double support in walking is primarily influenced by gait speed (27), and it is notable that in the current study, peak force beneath the heel was comparable with that reported elsewhere during steady-state walking in healthy adults (41). Moreover, heel pad properties have been shown to be relatively insensitive to physiological variations in loading rate (5,7). Finally, this study evaluated the force–deformation properties of the heel pad in healthy middle-age adults while walking at their preferred gait speed. Although previous research has reported an increase in heel pad stiffness with aging (19), such measures are confounded by the effects of tissue thickness, which have not been uniformly considered. Geometry-independent measures such as the elastic modulus or hysteretic energy, in contrast, have identified conflicting effects, with the EDR reported to be both significantly greater (3%–5%) and lower (12%) in the elderly (>60 yr) when compared with mature adults (18–40 yr) (14,16). Consequently, force–deformation properties reported for the heel pad in the current study may not be transferable to younger or older populations. Further research is required to determine the effects of maturation and aging on heel pad properties and to identify relative importance of gait kinematics, rearfoot motion, and relative soft tissue movement in the dissipation of impacts generated at speeds faster than preferred walking.
Transient loading profiles associated with barefoot walking induce rapidly changing deformation rates in the heel fat pad, yielding irregular force–deformation curves that are characterized by a point of inflection of the loading curve (at approximately 250 N) beyond which heel pad stiffness rises dramatically. Consistent with observations concerning the elastic modulus of the macro- and microchambers of the heel pad, the initial stiffness (32 N·mm−1) of the heel pad was 10 times lower than its final stiffness (211 N·mm−1), suggesting the heel pad may be viewed as a two-component series elastic structure. Although the energy-dissipating properties of the heel pad at physiologically relevant strain rates fall between those commonly cited for mechanical tests of cadaveric heels and impact loading in vivo, the peak deformation of the fat pad (10.3 mm) approached that predicted for the limit of pain tolerance, suggesting that the heel pad may operate close to its physiological maximum even at relatively moderate, preferred walking speeds.
The authors thank Dr. Dirk DeClerq for his assistance in the preparation of this manuscript.
The authors declare no conflicts of interest, financial or otherwise.
The results of the present study do not constitute endorsement by the American College of Sports Medicine.
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