As the foot is loaded with the weight of the body during the first interval, the calcaneus rapidly everts and the longitudinal arch flattens. This flattening of the arch originates in the subtalar joint and reaches a maximum during this interval (Fig. 1-2D). The hindfoot is often mildly supinated at initial ground contact associated with ankle dorsiflexion during swing-through. The hindfoot moving from supination to pronation during the first interval is a passive mechanism, and the amount of motion appears to depend entirely on the configuration of the articulating surfaces, their capsular attachments, and ligamentous support. No significant muscle function appears to play a role in restricting this motion at initial ground contact.

The subtalar joint links rotation of the hindfoot to rotation of the leg. During the first interval, eversion of the calcaneus is translated by the subtalar joint into inward rotation that is transmitted proximally across the ankle joint into the lower extremity (Fig. 1-2E). Distally, this hindfoot eversion unlocks the transverse tarsal joint (Fig. 1-2D), allowing the midfoot joints to become supple. This allows the flattening of the longitudinal arch that contributes to energy dissipation during this phase.

At heel strike, the center of gravity of the body is decelerated by ground contact, then immediately accelerated upward to carry it over the extending lower extremity. The heel’s impact and body’s center of gravity shift accounts for a vertical floor reaction that exceeds body weight by 15% to 25% (Fig. 1-2A).

Eccentric contraction of the anterior compartment leg muscles slows the rapid ankle plantar flexion during this phase from heel strike until a foot-flat position is reached. The posterior calf muscles all are electrically quiet, as are the intrinsic muscles in the sole of the foot (Fig. 1-2C). There is no muscular response in those muscles usually considered important in supporting the longitudinal arch of the foot. Weakness of the anterior compartment muscles leads to a loss of this deceleration and a characteristic slap foot gait.

Second Interval

The second interval extends from 15% to 40% of the walking cycle. During this interval, the body’s center of gravity passes from behind to in front of the weight-bearing leg. It reaches a maximum height as it passes over the leg at about 35% of the cycle, after which it commences to fall. During this interval, the foot transitions from a flexible, energy-absorbing structure to a more rigid one, capable of bearing the body’s weight.

The ankle joint undergoes progressive dorsiflexion during the second interval, reaching its peak at 40% of the walking cycle. This is when the force across the ankle joint has reached a maximum of 4.5 times body weight. Heel rise begins at 34% of the cycle as the contralateral leg passes by the stance foot and precedes the onset of plantar flexion, which begins at 40% (Fig. 1-3B).

Figure 1-3 Composite of events of second interval of walking, or period of foot flat. EMG, electromyograph.

During the second interval, the subtalar joint progressively inverts. This starts at about 30% of the cycle in a normal foot and at about 15% of the cycle in a flatfoot (Fig. 1-3D). Multiple factors contribute to this inversion, but precisely which plays the greatest role is unclear. Above the subtalar joint, the swinging contralateral limb externally rotates the stance limb. This external rotation torque is translated by the subtalar joint into hindfoot inversion. The oblique nature of the ankle joint axis, the oblique setting of the metatarsal break, and the function of the plantar aponeurosis also contribute to hindfoot inversion. Inversion of the subtalar joint is passed distally into the midfoot, increasing the stability of the transverse tarsal articulation and transforming the flexible midfoot into a rigid structure.

During this interval, full body weight is not borne on the foot, smoothing the transition to single limb support. Force plate recordings show that the load on the foot may be as low as 70% to 80% of actual body weight (Fig. 1-3A).

During the second interval, important functional changes occur in both the foot and leg, which are the result of muscular action. The posterior and lateral compartment leg muscles (triceps surae, peroneals, tibialis posterior, long toe flexors) and intrinsic muscles in the sole of the foot demonstrate electrical activity (Fig. 1-3C). Intrinsic muscle activity of the normal foot begins at 30% of the cycle, whereas in flatfoot, activity begins at 15% of the cycle. The posterior calf musculature slows the forward movement of the tibia over the fixed foot, which permits the contralateral limb to increase its step length. Weakness of the posterior compartment muscles may lead to premature contralateral heel strike and shortened stride length.

Third Interval

The third interval constitutes the last of the stance phase and extends from 40% to 62% of the walking cycle.

The ankle joint demonstrates rapid plantar flexion during this interval as the foot essentially extends the stance, effective length. The subtalar joint continues to invert during this interval, reaching its maximum at toe-off (Fig. 1-4D). This completes the conversion of the forefoot from the flexible structure observed in the first interval at the time of weight acceptance to a rigid structure at the end of the third interval in preparation for toe-off. The inversion is a continuation of the processes that began in the second interval. These include external rotation of limb above the foot passing across the ankle and subtalar joints as well as mechanisms in the foot such as the obliquity of the ankle joint, the function of the plantar aponeurosis, and obliquity of the metatarsal break. Distally, the transverse tarsal joint is converted from a flexible structure into a rigid one by the progressive inversion of the calcaneus. The talonavicular joint also is stabilized during this period by the pressure placed across the joint by both body weight and the intrinsic force created by the plantar aponeurosis.

Figure 1-4 Composite of all events of third interval of walking, or period extending from foot flat to toe-off. EMG, electromyograph.

At the beginning of the third interval, force plate recordings demonstrate an increase in the percentage of body weight borne by the foot resulting from the center of gravity falling. The load on the foot exceeds body weight by approximately 20%. Later in the interval, the vertical floor reaction force falls to zero as the body’s weight is transferred to the opposite foot (Fig. 1-4A).

Ankle plantar flexion during the third interval is caused primarily by the concentric contraction of the posterior calf musculature, in particular the triceps surae (Fig. 1-4B). The plantar flexion leads to relative elongation of the extremity. Although full plantar flexion at the ankle joint occurs during this interval, electrical activity is observed only until 50% of the cycle, after which there is no longer electrical activity in the extrinsic muscles (Fig. 1-4C). The remainder of ankle joint plantar flexion occurs because of the transfer of weight from the stance leg to the contralateral limb. The intrinsic muscles of the foot are active until toe-off. Although the intrinsic muscles help to stabilize the longitudinal arch, the main stabilizer is the plantar aponeurosis, which is functioning maximally during this period as the toes are brought into dorsiflexion and the plantar aponeurosis is wrapped around the metatarsal heads, forcing them into plantar flexion and elevating the longitudinal arch. The anterior compartment muscles become active in the last 5% of this interval, probably to initiate dorsiflexion of the ankle joint immediately after toe-off.

Running Cycle

The changes that occur in the gait cycle during running relative to walking are illustrated in Figure 1-5. During walking, one foot is always in contact with the ground; as the speed of gait increases, a transition occurs wherein a float phase is incorporated, during which time both feet are off the ground. Rather than a period of double limb support as occurs during walking, there is a period of no limb support. As the speed of gait continues to increase, the time the foot spends on the ground, both in real time and in percentage of cycle, decreases considerably. The speed at which one transitions from walking to running is greater than the speed at which one transitions back from running to walking.

Figure 1-5 Variations in gait cycle for running, jogging, and walking. Note that as the speed of gait increases, stance phase decreases. In this illustration, subject is walking at 3.75 miles per hour, jogging at about 1 mile per 9 minutes, and running about 1 mile per 5 minutes.

Kinematics of Human Locomotion

Humans use a unique and characteristic orthograde bipedal mode of locomotion. But walking is more than merely placing one foot in front the other. During walking, all major segments of the body are in motion. Displacements of the body segments occur in a well-preserved fashion and can be accurately described. Kinematics is the study of the motion of these body segments.

Human locomotion is a learned process; it does not develop as the result of an inborn reflex.⁵⁹ The first few steps of an infant holding onto his or her parent’s hand exemplify the learning process necessary to achieve orthograde progression. The result of this learning process is the integration of the neuromusculoskeletal mechanisms, with their gross similarities and individual variations, into an adequately functioning system of locomotion. Once a person has learned to walk, the mechanisms of ambulation are adaptable and work whether the person is an amputee learning to use a prosthesis, a long-distance runner, or a high-heeled shoe wearer.

A smoothly performing locomotor system results from the harmonious integration of many components. Because human locomotion involves all major segments of the body, certain suprapedal movements demand specific functions from the foot, and the manner in which the foot functions or fails to function may be reflected in patterns of movement in the other segments of the body. Similarly, alterations in movements above, such as a stiff knee or hip from arthritis or knee hyperextension from postpolio quadriceps weakness, may be reflected below by changes in the behavior of the foot.

Although bipedal locomotion imposes gross similarities in the manner in which all of us walk, each of us exhibits minor individual differences that allow us to be recognized by a friend or acquaintance, even from a distance. The causes of these individual characteristics of locomotion are many. Each of us differs somewhat in the length and distribution of mass of the various segments of the body, segments that must be moved by muscles of varying fiber length. Furthermore, individual differences occur in the position of axes of movement of the joints, with concomitant variations in effective lever arms. These and many more such factors combine to establish in each of us a final idiosyncratic manner of locomotion.

Just as no two people walk exactly alike, gait kinematics will not always be identical even within the same individual. The contribution of a single component varies under different circumstances. Type of shoe, amount of fatigue, weight of load carried, and other such variables can cause diminished functioning of some components, with compensatory increased functioning of others. An enormous number of variations in the behavior of individual components are possible; however, the diversely functioning components, when integrated, are complementary and will produce smooth forward progression.

Average values of single anthropometric observations of gait kinematic parameters are alone of little value. The surgeon should be alert to the anthropometric variations that occur within the population, but it is more important to understand the functional interrelationships among the various components. This is particularly true in the case of the foot, where anatomic variations are extensive. If average values are the only bases of comparison, it becomes difficult to explain why some feet function adequately and asymptomatically, although their measurements deviate from the average, whereas others function symptomatically, even though their measurements approximate the average. Therefore, in this chapter, emphasis is placed on functional interrelationships and not on lists of kinematic measurements.

At heel strike, progressive inward rotation occurs in the lower extremity, which consists of the pelvis, femur, and tibia, and this inward rotation reaches a maximum at the time of foot flat. The internal rotation at heel strike is initiated by the collapse of the subtalar joint into valgus, and its magnitude is determined by the flexibility of the foot and its ligamentous support. After contralateral toe-off, at about 12% of the cycle, progressive outward rotation occurs, which reaches a maximum at the time of toe-off, when inward rotation resumes (Fig. 1-8). Once the foot is on the ground, progressive external rotation is probably initiated by the contralateral swinging limb, which rotates the pelvis forward, imparting a certain degree of external rotation to the stance limb. This external rotation subsequently is passed from the pelvis distally to the femur and tibia, across the ankle joint, and is translated by the subtalar joint into inversion, which reaches its maximum at toe-off. The external rotation is enhanced by the external rotation of the ankle joint axis, the oblique metatarsal break, and the plantar aponeurosis after heel rise begins.

Figure 1-8 Transverse rotation occurring in the lower extremity during walking. Internal rotation occurs until approximately 15% of cycle, at which time progressive external rotation occurs until toe-off, when internal rotation begins again.

Kinetics of Human Locomotion

To begin a review of gait kinetics, one must recognize that the ambulating human is both a physical machine and a biologic organism subject to physical laws and beholden to muscular action. Gait kinematics and lower-extremity anatomic interrelationships strive to achieve a system that takes us from one spot to another with the least expenditure of energy.⁶⁰ Said another way, human locomotion is a blending of physical and biologic forces that combine to achieve maximum efficiency at minimum cost. Kinetics is the study of these energy expenditures.

All characteristics of muscular behavior are exploited in locomotion. Muscle groups may accelerate or decelerate body segments at different points in the gait cycle. They may contract concentrically (as they shorten) or eccentrically (as they lengthen). Part of energy conservation during the gait cycle involves having muscles work near their peak efficiency, which tends to be at or longer than their resting length.14,17,65 When motion in the skeletal segments is decelerated or when external forces work on the body, activated muscles become efficient. Activated muscles, in fact, are approximately six times as efficient when resisting elongation (eccentric contraction) as when shortening to perform external work.^1,5,6 In addition, noncontractile elements in muscles and specific connective tissue structures assist muscular action by providing an elastic component that stores and later releases kinetic energy.

Assessment of the forces and torques imparted by the ground on the lower extremity has illuminated the biomechanical processes at work during gait. Investigation of the pressures experienced by the various regions of the plantar foot has provided insight into the pathogenesis and treatment of many foot and ankle disorders. A number of tools have evolved to study gait kinetics. These are described in detail in the next section, followed by an analysis of kinetics during gait.

Measuring Whole Body Kinetics and Plantar Pressure

Studying the foot’s interaction with the ground has a long history, ranging from examining footprints in soil to real-time mapping of plantar pressure under natural conditions. Plantar pressure and ground reaction force measurements are well established in the research realm and have been instrumental in refining our understanding of foot and ankle biomechanics. In conjunction with other technology, including high-speed cameras, video motion-sensing equipment, electrogoniometers, and electromyograph (EMG) devices, the study of the ground–foot interaction has aided the understanding of gait kinetics and kinematics.

Despite improvements in available measurement methods, however, practical collection of clinically novel information remains difficult. The wide variability of normal measures makes clinical comparisons difficult. The large number of measurement systems and equally large number of data analysis techniques make it difficult to generalize results.

Although confirmation of areas of excess pressure and monitoring the effects of treatment may prove useful, there is little specificity between plantar pressure patterns and clinical syndromes.

Types of Studies

A variety of measurement techniques have been used to study the interaction of the foot with the ground. Indirect techniques rely on correlating other measurable gait parameters to plantar characteristics and offer the advantage of not relying on expensive and often bulky equipment. For example, an estimation of ground reaction force can be made based on a simple-to-measure temporal variable, foot–ground contact time.¹³

Direct measurement techniques rely on physical properties or electronic transducers to translate the interaction between the foot and the ground into a measurable quantity. Multiple direct measurement systems are available that use a variety of strategies to record plantar pressure or ground reaction force. Unfortunately, results obtained with different systems under similar conditions are not always similar, and even qualitative comparisons may not be appropriate.³⁸ Spatial resolution and sample rate affect the ability of a system to record true peak plantar pressures and to isolate particular areas under the foot.

The earliest direct measurement methods relied on physical properties of a material to capture the interaction of the foot with the ground. Casts of the foot in clay, plaster, or soil were used with the assumption that areas of deeper penetration represented areas of highest pressure.10,21 Rubber mats incorporating longitudinal ridges,⁵⁴ pyramidal projections,²¹ or a multilevel grid (such as the Harris-Beath mat),^67,78 use the elastic property of rubber which, when stood or walked on, distorts in proportion to the pressure applied (Fig. 1-9). Although fast, inexpensive, and portable, these methods have low measurement resolution and lack temporal discrimination.⁶⁷

Figure 1-9 Pressure distribution on plantar aspect of foot as demonstrated by use of barograph. As dots get larger and denser, pressure distribution is greater. (From Elftman H: A cinematic study of the distribution of pressure in the human foot. Anat Rec 59:481-491, 1934.)

Optically based systems rely on visualizing the plantar aspect of the foot during stance or gait. The simplest allows observation or photographic recording of the plantar foot through a clear platform (Fig. 1-10). This provides an accurate, dynamic, qualitative representation of foot morphology. Addition of a physical transduction device between the foot and glass plate allows quantification of regionalized plantar pressures and adds the temporal component missed using a physical transduction system alone.²¹ The pedobarograph places a thin plastic sheet over the clear plate.⁴ The sheet is illuminated at the edges, and pressure on the plastic distorts the light in proportion to the pressure applied. The images can be recorded and calibrated to provide a spatial resolution and temporal responsiveness not found with the Harris-Beath mat. However, slow responsiveness at high forces may bias results.³⁶

Figure 1-10 Feet and legs of person standing on barograph. A, Weight bearing with muscles relaxed. B, Rising on toes.

A force plate measures the ground reaction force, that is, the force exerted by the ground on the foot, in three degrees of freedom (vertical force, forward shear, side shear), and allows calculation of the torques around the foot and ankle (axial torque, sagittal torque, coronal torque). Force transducers are configured in orthogonal planes at the corners of a section of floor. The resulting data provide a representation of the average forces experienced by the foot over the gait cycle (Fig. 1-11). One advantage of this type of system is that shear forces and torques can be measured in addition to vertical force. The limitations include the lack of ability to map specific regions of plantar pressure. This limitation can be circumvented with the addition of an optical diffraction system, as described above, or with a series of smaller force plates placed in tandem.⁷¹

Figure 1-11 Ground reaction to walking. A, Vertical force. B, Fore–aft shear. C, Medial–lateral shear. D, Torque. Ext., External; Fwd., forward; Int., internal; Lat., lateral; Med., medial.

The ability to place pressure transducers on discrete parts of the foot has become possible, as their size has shrunk. They can be placed on strategic points of the foot, or an array can be created to map the pressures exerted by the foot during stance or gait. These data provide a spatial and temporal map of plantar pressure over the gait cycle 54,62 (Fig. 1-12). Many of these systems use a floor mat or platform built into the floor with a grid of pressure-sensitive transducers. An alternative is to place a thin film containing a pressure transduction array into a shoe (Fig. 1-13). In this way, the plantar pressures experienced by the foot can be measured in a wider variety of settings and under multiple impacts as well as account for the effect of shoe wear.^42,77 For example, feet experience 10% to 50% higher plantar pressures in a flat, flexible shoe compared with a soft shoe with a firm rubber sole.⁶³ The floor mat and in-shoe methods correlate well when the shoe used has a firm sole or when barefoot.⁷

Figure 1-12 Pressure distribution under bare foot during walking. Height of display above ground is proportional to pressure. (From Clarke TE: The pressure distribution under the foot during barefoot walking [doctoral dissertation], University Park, Pa, 1980, Pennsylvania State University.)

Figure 1-13 Peak plantar pressure map using an in-shoe thin-film pressure transducer. Red represents areas of relatively high pressure, and violet, areas of low pressure. (Courtesy Ken Hunt, MD.)

A number of system-specific and analysis-dependent factors affect the results of pressure transducer array measurements, including pressure transducer density, responsiveness, linearity, resolution, and range of the transducers. Methods of analyzing the data also differ, including reporting results as force versus pressure, peak values versus sum of values over time, and strategies of regionalizing the foot’s plantar surface. Increasing pressure transducer density provides better spatial representation of plantar pressure, whereas systems with relatively lower transducer density may underestimate measurements, such as peak pressure, because the true peak may be missed. Some transducers may have a nonlinear response at the extremes of their measurable range or have a low-level cutoff. The maximum sample rate affects contact time measurements, and low sample rates may underestimate peak pressure measurements because the true peak pressure may be missed.

Data Representations

Output from the different measurement systems reflects the nature of their measurement mechanisms. The Harris mat reports pedal pressure but does not vary with time. The force plate reports a true ground reaction force but in not spatially discriminative. The optical systems and the transduction arrays each report pedal pressure that varies with time. The data measured by these systems is subject to sensor density, resolution, and sample rate limitations discussed above. To simplify the information and allow comparisons between subjects or after treatments, a variety of derivative parameters have been defined based on these raw data. Not all systems or measurement methods are able to derive all of these measurements.

The ground reaction force is a vector quantity varying temporally and spatially over the gait cycle that represents the average reciprocal force exerted by the floor in response to the foot. It has a magnitude and direction, and the starting point may be projected onto a representation of the plantar foot at the point of average maximum vertical force (Fig. 1-14). The ground reaction force can be deconstructed into vertical force, anterior–posterior shear, and medial–lateral shear. The vertical ground reaction force represents the force of the ground pushing upward on the foot, and can be calculated from systems that measure plantar pressure for the whole foot or for defined regions of the foot.⁷⁶ Typically, it has two peaks; the first peak occurs as the body weight is transferred from dual- to single-leg stance and the second as the body weight moves forward over the metatarsal heads. Studies of ground reaction forces may focus on the magnitude of one or the other vertical peak or the timing of the peaks and valleys. Torque (moment) and power around a joint also can be calculated from the ground reaction force, joint geometry, timing parameters, and kinematics.

Figure 1-14 Peak pressure values under selected foot regions demonstrate impact in heel region, minimal weight bearing in midfoot, buildup of pressure beneath metatarsal heads, and transfer of weight to great toe region. (From Hennig EM, Rosenbaum D: Pressure distribution patterns under the feet of children in comparison with adults. Foot Ankle 11:306-311, 1991.)

Another frequently reported measurement is the maximum pressure recorded, or peak pressure. It is usually reported over a spatially subdivided map of the plantar foot. Peak pressure for areas such as the heel, individual or grouped metatarsal heads, and toes are common. Alternatively, peak pressure can be reported as a temporally varying measure by displaying its location and magnitude on a diagram of the foot. Peak force can be calculated from peak pressure because the size of the pressure transducers is known. Calculated joint moments represent the torque applied by muscles to counteract the measured ground reaction force, and joint power is calculated from the joint moment and angular velocity (Fig. 1-15).

Figure 1-15 Joint power generated (positive) or absorbed (negative) during the gait cycle before and after total ankle replacement. (Modified from Brodsky JW, Polo FE, Coleman SC, Bruck N: Changes in gait following the Scandinavian Total Ankle Replacement. J Bone Joint Surg Am 93:1890-1896, 2011.)

Timing measurements can also be made. The time intervals from heel strike to metatarsal strike, toe strike, heel-off, metatarsal-off, and toe-off can be calculated. The pressure·time integral, or impulse, for the whole foot or defined regions can be calculated. This may be standardized for each region as a percentage of the total impulse for a given foot. The impulse may characterize plantar loading better than peak pressure by taking both pressure and time into consideration.

Finally, the pattern of plantar loading can be categorized based on the pressure measurements. Patients may tend to load the medial ray, the medial and central rays, the central rays, or the central and lateral rays.³⁵ Put another way, there is an inverse relationship between peak pressure under the first metatarsal head and toe relative to the lesser metatarsal heads.²⁵ As walking speed increases, a medialization of forefoot pressure occurs such that peak pressure under the first metatarsal head increases and that under the lesser metatarsal heads decreases.⁶⁴

Measurement Variability

Many sources of variability affect the results of these measurements. Separating important clinical or research findings from differences based on testing apparatus, measurement methodology, patient demographic factors, or analysis methodology requires an understanding of how these factors affect the measured results. Differences between the different testing apparatus have been described above. Other sources of variability can be divided into methodology, analysis, and patient-specific factors.

Walking speed affects the magnitude of plantar pressures during gait. Velocity is linearly related to peak vertical and fore–aft ground reaction forces,3,57 and inversely related to the pressure·time integral.⁸¹ As velocity increases, peak pressures on the heel, medial metatarsal heads, and the first toe increase while peak pressure in the fifth metatarsal head decreases.^35,64 This medialization may be related to increased magnitude and velocity of hindfoot eversion and medial shear force at heel strike. Timing measurements also change with increasing speed. The normalized time to peak pressure is decreased on the heel but unchanged in the midfoot and forefoot, suggesting the rollover process is mainly accelerated by reducing the time from heel strike to foot flat.⁶⁴ To minimize variability introduced because of walking speed, subjects may walk at a fixed rate or at their natural pace.⁸¹

Deviations from a normal gait pattern can occur if the subject has to take a long or short stride in an effort to place the foot on the appropriate measurement area of floor-based systems. To minimize this effect, the measurement platform is placed flush with the floor and hidden from the subject with a thin, uniform floor covering. The traditional midgait method uses a short lead-up walk before the foot strikes the measurement platform. A three-step or two-step lead-up is as reproducible, but a one-step lead-up is not adequate.15,56

Variability of the measurements is also dependent on the type of gait. For example, plantar pressures measured when standing differ from pressures measured during gait.¹⁰ Variations in walking patterns, such as a shuffling-type gait, alter the peak forces on the foot.⁸² Gait pattern alteration can be seen in certain conditions, such as after ankle fracture fixation or with concurrent knee pathology.^3,8

Drift and calibration of the measurement systems affect the variability of measurements. Plantar pressure measurement systems need to be calibrated to allow comparisons between systems. Transducer output varies between different transducers, with temperature, when an in-shoe system is removed and reinserted, and with the number of trials performed. Pressure can vary by as much as 20% with repeated measurements on the same insert.⁶³ There may be an offset that drifts with time.⁵⁶ The measurements may be adequate for relative ranking purposes but need repeated calibration with a fixed system if accurate values are needed.

Variability is also introduced in the methods by which the acquired data are analyzed. For example, peak pressure can be reported for the whole surface of the foot during a gait cycle, but the clinical utility of this is limited because different regions of the foot experience different plantar pressures during the gait cycle. Subdividing the regions of the plantar foot and recording peak pressures in each of these areas over the gait cycle provides more meaningful data. The heel is often represented as a single region but may be subdivided into medial, central, and lateral.⁶⁴ Midfoot peak pressures may be useful in pathologic conditions, such as rocker-bottom deformity, and can classify foot morphology into planus, normal, and cavus categories.⁶⁴ The base of the fifth metatarsal can be included as part of the midfoot or can be identified as a separate pressure zone. Improvements in sensor technology have allowed measurement of individual metatarsal heads and toe forefoot pressures.^25,37 Definition of these regions (masks) is still a manual process and is repeated for each trial. Having a single person define the regions may decrease variability.³⁷

Subject-specific characteristics also introduce variability. Children’s feet have a dramatically different loading pattern and lower peak pressure because of high relative foot area ²⁷ (see Fig. 1-14). Differences in joint mobility and forefoot pressure based on a subject’s ethnicity have been shown in neuropathic diabetics.⁷³ The patient’s dominant side may experience greater static and dynamic vertical force,⁴² although others have found no side dominance.²⁹ Foot morphology also affects plantar pressure; cavus feet have different midfoot loading characteristics and rate and degree of hindfoot eversion than flatfeet.⁶⁴ During running, fatigued subjects tend to have decreased step time, decreased peak and integral force and pressure under heel, and medialization of forces.⁷⁷ After a hindfoot fusion, greater contact force at heel strike has been observed.^45,69 This could be due to the inability of the calcaneus to move into a valgus position after heel strike.²

The effect of body weight on plantar pressure is less direct than might be expected. Although some have correlated maximum vertical force during gait and body weight,⁴² many other studies found little correlation.^26,27,39 In children, the correlation of body weight to peak plantar pressures is clear and plays a greater role in determining peak pressure than in adults.^27,35 The area of peak pressure most highly correlated with body weight in children is the fourth metatarsal head,³⁵ and in adults may be the fourth metatarsal head or the midfoot.^27,35