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Normal Human Locomotion, Part 1: Basic Concepts and Terminology Ed Ayyappa, MS, CPO

ABSTRACT

Over the past several decades, the evolution of gait science has produced an array of terms and concepts relating to observational gait analysis. Prosthetists and orthotists use various forms of gait analysis on a daily basis as an important part of clinical care.

When the basic principles of normal walking are understood, a more penetrating grasp of pathological gait becomes possible. The result is expanded ability to differentiate between pathological and compensatory gait deficits. In addition, efficient interaction with the clinic team demands a sound conceptual knowledge base of human locomotion and related terminology. This will facilitate an optimal treatment plan for the patient and enhance communication and prescription recommendations to the physician and relevant paying agencies. This article is intended to be an introduction to gait science with these goals and objectives in mind.

Introduction

Nearly a century ago, A.A. Marks, an American prosthetist, offered a precise qualitative description of normal human locomotion when he illustrated and analyzed the walking process in eight organized phases and discussed the implications of prosthetic design on the function of amputee gait (see Figure 1) . Well ahead of his time, Marks praised "kinetoscopic" photography as a potential diagnostic tool for the improvement of walking deficits (1).

Insight into normal walking patterns can help practitioners improve the efficiency of persons with gait-related pathologies. Such knowledge may assist the clinician in the selection of orthotic or prosthetic componentry, alignment parameters and identification of other variants that may enhance performance (2). Familiarity with gait terminology and function enables the prosthetist or orthotist to communicate effectively with other members of the medical team and contributes to the development of a sound treatment plan.

The terminology of human walking began with descriptive phrases obtained as a result of observational and kinematic analysis of normal subjects. This approach yielded terms such as "push off" and "heel strike" (as differentiated from "foot flat"). The limitations of these terms for clinical use became apparent as practitioners' understanding of normal locomotion increased and was melded with a careful observation of pathological function. "Push off," for example, is a misleading term because in free-walk velocity during the last period of stance phase (preswing), the posterior compartment musculature is quiescent. While a differentiated heel strike and foot flat may describe normal function, these terms are woefully inadequate in describing the common clinical picture of an equinus stance phase. Many more contemporary terms describe events and functions that were not apparent through observation but could be measured through instrumentation in gait laboratories.

The separate contributions of Saunders et al. (3), Perry (4), Sutherland (5,6) and others have increased practitioners' understanding of gait science and terminology. Decades of work by Jacquelin Perry, MD, have resulted in descriptive terms for the phases and functional tasks of gait (7). These phases and tasks have received wide acceptance and serve as the descriptive medium for this article. Contemporary terminology continues to evolve through dialogue within professional organizations such as the North American Society of Gait and Clinical Movement Analysis (8) and the American Academy of Orthotists and Prosthetists (AAOP) Gait Society (9). This article, presented as an introduction to the AAOP Professional Development Certificate Program in Gait and Pathomechanics (10), attempts to reflect current contemporary usage of gait terminology.

Divisions of Gait Cycle

Gait characteristics are influenced by the shape, position and function of neuromuscular and musculoskeletal structures as well as by the ligamentous and capsular constraints of the joints. The primary goal is energy efficiency in progression using a stable kinetic chain of joints and limb segments that work congruently to transport the passenger unit-head, arms and trunk (HAT). The lower extremities and pelvis, which carry the HAT, are referred to as the locomotor apparatus.

The gait cycle is the period of time between any two identical events in the walking cycle. Any event could be selected as the onset of the gait cycle because the various events follow each other continuously and smoothly. Initial contact, however, generally has been selected as the starting and completing event.

By contrast, the gait stride is the distance from initial contact of one foot to the following initial contact of the same foot.

Each gait cycle is divided into two periods, stance and swing. Stance is the time when the foot is in contact with the ground, constituting 62 percent of the gait cycle. Swing denotes the time when the foot is in the air, constituting the remaining 38 percent of the gait cycle. In those cases where the foot never leaves the ground, sometimes referred to as foot drag, the swing phase could be defined as the phase when all portions of the foot are in forward motion.

Double support is the period of time when both feet are in contact with the ground. This occurs twice in the gait cycle-at the beginning and end of stance phase-and also is referred to as initial and terminal double-limb stance (see Figure 2) . As velocity increases, double-limb support time decreases. Running constitutes forward movement with no period of double-limb support.

In normal walking, initial double-limb support takes up about 12 percent of the gait cycle, and terminal double-limb support occupies 12 percent as well. Generally, the two periods of double-limb support represent 25 percent of the gait cycle.

Single support is the period of time when only one foot is in contact with the ground. In walking, this is equal to the swing phase of the other limb. The term ipsilateral is used to describe the same side of the body, and the term contralateral is used to describe the opposite side of the body or the opposite limb. The direction of walking is referred to as the line of progression.

Functional Tasks of Gait

A complete gait cycle can be viewed in terms of three functional tasks of weight acceptance, single-limb support and limb advancement (see Figure 3) .

The first functional task is weight acceptance. Two phases of the stance period, initial contact and loading response, are involved in the performance of weight acceptance. The demand for immediate transfer of body weight onto the limb as soon as it contacts the ground requires initial limb stability and shock absorption while simultaneously preserving the momentum of progression. When the functional task of weight acceptance has been achieved, the individual is said to demonstrate a stable kinetic chain.

The second functional task is single- limb support. Primarily, two phases are associated with single-limb support: midstance and terminal stance. In addition, preswing is a transitional phase that could be considered part of single-limb support as well. During this period, the contralateral foot is in the swing period, and total body weight is exclusively supported on the stance limb. Forward progression of body weight over the stationary foot while maintaining stability is accomplished.

The third functional task is limb advancement. Four phases contribute to limb advancement: preswing, initial swing, midswing and terminal swing. During these phases, the stance limb leaves the ground and advances forward to posture itself in preparation for the next initial contact. The preswing phase serves in both single-limb support and limb advancement.

Phases of Gait

The gait cycle can be described in the phasic terms of initial contact (IC), loading response (LR), midstance (MSt), terminal stance (TSt), preswing (PSw), initial swing (ISw), midswing (MSw) and terminal swing (TSw) (see Figure 3) . The stance period consists of the first five phases: initial contact, loading response, midstance, terminal stance and preswing. The swing period primarily is divided into three phases: initial swing, midswing and terminal swing. Preswing, however, prepares the limb for swing advancement and in that sense could be considered a component of swing phase.

Initial Contact

Initial contact is an instantaneous point in time only and occurs the instant the foot of the leading lower limb touches the ground. Most of the motor function that occurs during initial contact is in preparation for the loading response phase that will follow.

Initial contact represents the beginning of the stance phase. Heel strike and heel contact serve as poor descriptors of this period since there are many circumstances when initial contact is not made with the heel alone. The term "foot strike" sometimes is used as an alternative descriptor.

Loading Response

The loading response phase occupies about 10 percent of the gait cycle and constitutes the period of initial double-limb support. During loading response, the foot comes in full contact with the floor, and body weight is fully transferred onto the stance limb.

The initial double-support stance period occasionally is referred to as initial stance. The term foot flat (FF) is the point in time when the foot becomes plantar grade. The loading response period probably is best described by the typical quantified values of the vertical force curve. The ascending initial peak of the vertical force graph reveals the period of loading response (see Figure 4) .

Midstance

Midstance represents the first half of single support, which occurs from the 10- to 30-percent periods of the gait cycle. It begins when the contralateral foot leaves the ground and continues as the body weight travels along the length of the foot until it is aligned over the forefoot. The descending initial peak of the vertical force graph reveals the period of midstance (see Figure 4) .

Terminal Stance

Terminal stance constitutes the second half of single-limb support. It begins with heel rise and ends when the contralateral foot contacts the ground. Terminal stance occurs from the 30- to 50- percent periods of the gait cycle. During this phase, body weight moves ahead of the forefoot.

The term heel off (HO) is a descriptor useful in observational analysis and is the point during the stance phase when the heel leaves the ground. The ascending second peak of the vertical force graph demonstrates the period of terminal stance (see Figure 4) .

Roll off describes the period of late stance (from the 40- to 50- percent periods of the gait cycle) when there is an ankle plantarflexor moment and simultaneous power generation of the triceps surae to initiate advancement of the tibia over the fulcrum of the metatarsal heads in preparation for the next phase.

Preswing

Preswing is the terminal double-limb support period and occupies the last 12 percent of stance phase, from 50 percent to 62 percent. It begins when the contralateral foot contacts the ground and ends with ipsilateral toe off. During this period, the stance limb is unloaded and body weight is transferred onto the contralateral limb. The descending portion of the second peak of the vertical force graph demonstrates the period of preswing (see Figure 4) .

Terminal contact (TC), a term rarely used, describes the instantaneous point in the gait cycle when the foot leaves the ground. It thus represents either the end of the stance phase or the beginning of swing phase. In pathologies where the foot never leaves the ground, the term foot drag is used. In foot drag, the termination of stance and the onset of swing may be somewhat arbitrary.

The termination of stance and the onset of swing is defined as the point where all portions of the foot have achieved motion relative to the floor. Likewise, the termination of swing and the onset of stance may be defined as the point when the foot ends motion relative to the floor. Toe off occurs when terminal contact is made with the toe.

Initial Swing

The initial one-third of the swing period, from the 62- to 75-percent periods of the gait cycle (6), is spent in initial swing. It begins the moment the foot leaves the ground and continues until maximum knee flexion occurs, when the swinging extremity is directly under the body and directly opposite the stance limb.

Midswing

Midswing occurs in the second third of the swing period, from the 75- to 85-percent periods of the gait cycle (6). Critical events include continued limb advancement and foot clearance. This phase begins following maximum knee flexion and ends when the tibia is in a vertical position.

Terminal Swing

In the final phase of terminal swing from the 85- to 100-percent periods of the gait cycle (6), the tibia passes beyond perpendicular, and the knee fully extends in preparation for heel contact.

Temporal Parameters

The potential to assess gait through quantified measurement emerged with the sunrise-to-sunset movement of a lone traveler on foot over a known distance or with the hailing chant of each advancing step of a marching army. Such events would have enabled measurement of walking velocity, or distance traversed per unit of time, and cadence, or steps per unit of time. Gait parameters related to time are referred to as temporal parameters.

Stride length, cadence and velocity are three important interrelated temporal parameters. Commonly misused, the terms step length and stride length are not synonymous. As a dynamic measurement of gait, step length is the distance in meters from a given floor-contact point of the ipsilateral (or originating) foot in stance to the same floor-contact point of the contralateral (or opposite) foot in stance (see Figure 5) ; for example, the distance from right-heel contact to left-heel contact. The step period is the segment of time in seconds taken for one step to occur and is measured from an event of one foot to the following occurrence of the same event with the other foot.

Stride length, on the other hand, contains both a left- and a right-step length (see Figure 5) (e.g., the distance from right-heel contact to the following right-heel contact). Stride length sometimes is referred to as cycle length and is expressed in meters. A reduction in functional joint motion or the presence of pain or muscle weakness may reflect a reduction in stride or step length. Pathological gait commonly produces asymmetries in step length between limbs. Stride period or cycle time is the period of time in seconds from initial contact of one foot to the following initial contact of the same foot.

Cadence refers to the number of steps taken per unit of time and is the rate at which a person walks expressed in steps per minute. Natural or free cadence describes a self-selected walking rhythm.

Velocity combines stride length and cadence and is the resultant rate of forward progression. Velocity is the rate of change of linear displacement along the direction of progression

measured over one or more strides and is expressed in meters per second. It is the best single index of walking ability.

Reductions in velocity correlate with joint impairments, amputation levels and many acute pathologies. Velocity may be quantitatively measured or qualitatively assessed using the terms free, slow and fast. Free walking speed describes the normal self-selected walking velocity. Fast walking speed describes the maximum velocity attainable by a subject with a pathological gait. Slow walking speed describes a velocity below the normal self-selected walking speed. Fast walk velocity for healthy subjects can increase by as much as 44 percent (11), but pathological subjects have less buffer. Since velocity affects many parameters of walking, the typical description of normal gait generally presupposes a comfortable self-selected velocity. With this free walk velocity, the individual will naturally enlist both the mannerisms and speed that will provide maximum energy efficiency.

Temporal parameters historically have been obtained in a gait lab by means of microswitches embedded in plantar foot pads taped to the bottom of the foot or shoe (12) (see Figure 6) . The rollover patterns are recorded as the patient walks a measured distance, and the temporal parameters are calculated.

Although microswitches have been the standard for some time, perhaps the most promising measurement tools for collecting temporal data are pressure-sensing arrays. A thin plastic sheet of film can slip nearly unnoticed between the plantar surface of the foot and the orthosis within the shoe (see Figure 7) . This array, connected to a computer via a lead wire, can measure dynamic pressure patterns and record critical events throughout the walking cycle. A prosthetic version can provide pressure measurements at 60 individual sites within a socket and record those measurements during multiple events of the gait cycle. The current clinical relevance lies in identifying critical gait events and skin-loading pressure patterns. Because of the ease in collection of plantar pressure readings and relative modest cost, this approach may well replace microswitch technologies in the future and be increasingly accessible to prosthetists and orthotists for clinical use.

Microswitch technologies enable the clinician to record tendencies toward excessive inversion, eversion or prolonged heel-only time and can suggest modifications to alignment or componentry of prostheses or orthoses to normalize such patterns.

Time-distance parameters have enormous potential for setting outcome goals. Variations in time-distance values often are pathology-specific. Asymmetries in hemiplegics, for example, obviously are greater than in most other pathologies; this technology is uniquely suited for quantifying those asymmetries.

As basic temporal technologies develop and become increasingly affordable and as mean pathology-specific values are obtained, these time-distance parameters, captured from microswitch or piezoelectric film pressure technology, may become the baseline for measuring functional outcomes.

Determinants of Gait

Saunders et al. defined walking as the translation of the center of mass through space in a manner requiring the least energy expenditure. They identified six determinants or variables that affect that energy expenditure (3). Variations in pelvic rotation, pelvic tilt, knee flexion at

midstance, foot and ankle motion, knee motion, and lateral pelvic displacement all affect energy expenditure and the mechanical efficiency of walking.

As a functional basis for understanding energy efficiency in gait, these principles have stood the test of time (13-15). These determinants of gait are based on two principles: 1) Any displacement that elevates, depresses or moves the center of mass beyond normal maximum excursion limits wastes energy, and 2) Any abrupt or irregular movement will waste energy even when that movement does not exceed the normal maximum displacement limits of the center of mass. A successful long-distance runner intuitively takes advantage of these principles. By contrast, the unsuccessful runner lumbers from side to side and lurches up and down in a vicious spiral of exhaustion followed by increased energy expenditure.

Of the six determinants of gait, three provide mechanical advantages that limit vertical displacement of the center of mass. The term center of mass is synonymous with the term center of gravity (CG). Without these mechanical advantages that limit displacement, the center of mass would displace vertically 7.5 cm (3 inches) on a person of average height. Resulting from these three determinants, the center of mass is said to displace vertically only 5 cm (2 inches).

Pelvic Rotation

The trailing extended weight-bearing limb is elastically linked through the joints of the pelvis with the advancing swing limb. Ligamentous constraints and muscular activity combine with forward momentum of the advancing swing limb to position the pelvis into four degrees of rotation from the line of progression prior to initial contact (see Figure 8) . During the reciprocating contralateral swing phase, the pelvis rotates in the opposite direction, first returning to its neutral position and then continuing to rotate an additional 4 degrees. Thus the total range of pelvic rotation is 8 degrees.

In the act of pelvic rotation, both the trailing and advancing limbs are effectively lengthened through the rotation that uses the pelvic width to extend both support points. At the very time when the center of mass would otherwise drop excessively, this rotation prevents .95 cm (3/8-inch) of downward displacement of the center of mass.

Pelvic Tilt

At midstance, the center of mass reaches its highest point as the body vaults over a planted leg. It would be even higher were it not for the pelvis, which tilts down toward the swing side 5 degrees from vertical (positive Trendelenburg sign) and thus depresses the center of mass .5 cm (3/16-inch) in an efficient method of energy conservation. This is referred to as pelvic list or pelvic tilt and is possible only in conjunction with adequate limb clearance in swing phase (see Figure 9) .

Knee Flexion During Midstance

The stance limb enters initial contact with the knee in nearly full extension. It then flexes as the foot shifts to a plantar-grade position and continues moving into flexion until it reaches approximately 15 degrees. The knee then begins to extend but retains some flexion as it nears midstance; due to a relatively less extended knee as the tibia reaches verticality when the

center of mass is at its peak, the summit of the CG is depressed in its elevation by 1.1 cm (7/16-inch).

To summarize, the .95-cm displacement savings from pelvic rotation, .5-cm savings from pelvic tilt and 1.1-cm savings derived from knee flexion at midstance result in a combined displacement savings of 2.1 cm (approximately 1 inch). Without these three determinants, pelvic rotation, pelvic tilt and knee flexion at midstance, the vertical displacement of the center of mass is thought to be 7.5 cm (3 inches). With the 2.1 cm savings derived from these determinants, the vertical displacement of the center of mass has been reduced to approximately 5 cm (2 inches). If these three determinants were the only mechanisms affecting the progression of the center of mass as it traverses through space, the CG pathway would consist of a series of arcs at whose intersections an abrupt shift in the direction of the CG would occur as it reached its lowest point. However, both foot and ankle motion as well as knee motion serve to smooth the pathway of the CG.

Foot and Ankle Motion

The most important mechanism to smooth this pathway is foot and ankle motion. At initial contact, the ankle is elevated due to the heel lever arm but falls as the foot becomes plantar grade. At heel rise, the ankle again is elevated, which continues through terminal stance and preswing. These ankle motions, coordinated with the knee and controlled by muscle action of pretibials and triceps surae, smooth the pathway of the center of mass during stance phase (see Figure 11) .

The controlled lever arm of the forefoot at preswing is particularly helpful as it rounds out the sharp downward reversal of the center of mass. Thus it does not reduce a peak displacement period of the center of mass as the earlier determinants did but rather smooths the pathway. Foot and ankle motion thus facilitate the path of the CG, keeping it relatively horizontal throughout stance phase.

Knee Motion

Knee motion is intrinsically associated with foot and ankle motion. At initial contact before the ankle moves into a plantar-grade position and thus is relatively more elevated, the knee is in relative extension. Responding to a plan- tar-grade posture (when the ankle is depressed), the knee flexes. Passing through midstance as the ankle remains stationary with the foot flat on the floor, the knee again reverses its direction to one of extension. As the heel comes off the floor in terminal stance, the ankle again is elevated, and the knee flexes. In preswing, as the forefoot rolls over the metatarsal heads, the ankle moves even higher in elevation as flexion of the knee increases (see Figure 12) . Generally, at periods when the ankle center is depressed, the knee extends, and at periods when the ankle is elevated, the knee flexes. Knee motion, intimately associated with foot and ankle motion, smooths the pathway of the center of mass and thus conserves energy.

Lateral Pelvic Displacement

To avoid extraordinary muscular and balancing demands, the pelvis shifts over the support point of the stance limb. If the lower extremities dropped directly vertical from the hip joint, the center of mass would be required to shift three to four inches to each side to be positioned effectively over the supporting foot. The combination of femoral varus and anatomical

valgum at the knee permits a vertical tibial posture with both tibias in close proximity to each other. This narrows the walking base to 5-10 cm (2-4 inches) from heel center to heel center. This reduces the lateral shift required of the center of mass to 2.5 cm (1 inch) toward either side (see Figure 13) . The walking base or stride/step width typically is measured from one ankle joint center to the other although it often is described as the measurement from heel center to heel center.

A wide walking base may increase stability-but at a cost of energy efficiency-and the center of mass remains in a box two inches tall and two inches wide as the individual ambulates forward in normal human locomotion.

Foot and Ankle Function: The Rocker Mechanisms

Perry has described the function of the heel, ankle and forefoot rocker mechanisms in normal gait (4). Understanding the natural mechanics of these rockers greatly improves the abilities to diagnose and communicate orthotic and prosthetic gait deficits.

The first rocker is referred to as the heel rocker. The momentum generated by the fall of body weight onto the stance limb is preserved by this heel rocker. Normal initial contact is made by the calcaneal tuberosity, which becomes the fulcrum about which the foot and tibia move. The bony segment between this fulcrum and the center of the ankle rolls toward the ground as body weight is dropped onto the stance foot, preserving the momentum of forward progression.

The second rocker is the ankle rocker. The pivotal arc of the ankle rocker advances the tibia over the stationary foot.

The third rocker is referred to as the forefoot rocker. During terminal stance, as the body vector approaches the metatarsal-phalangeal joint, the heel rises and the phalanx extend. The metatarsal heads serve as an axis of rotation for body weight advancement.

The location of the ground-reaction force during preswing and concurrent loading on the contralateral limb enables passive knee flexion, which prepares the ipsilateral limb for the clearance demands of swing phase.

Conclusion

Normal bipedal gait is achieved with a complex combination of automatic and volitional postural components. Normal walking requires stability to provide antigravity support of body weight, mobility of body segments and motor control to sequence multiple segments while transferring body weight from one limb to another. The result is energy-efficient forward progression. (5)

Acknowledgements

The author would like to express appreciation to Ken Hudgens, program manager of the prosthetic and orthotic department, California State University - Dominguez Hills, for illustrations 4, 10, 11 and 12.

References:

1. Marks AA. Manual of artificial limbs. New York: A.A. Marks Inc., 1905:17-20. 2. Ayyappa E, Mohamed O. Orthotics and prosthetics in rehabilitation. In: Lusardi M,

ed. Clinical assessment of pathological gait. Newton, Mass.: Butterworth Heinemann, manuscript submitted for publication, September 1996.

3. Saunders JB, Inman VT, Eberhart HD. The major determinants in normal and pathological gait. JBJS 1953; 35-A:543-58.

4. Perry J. Gait analysis; normal and pathological function. Thorofare, N.J.: Slack, 1992. 5. Sutherland D. Development of mature walking. Philadelphia: MacKeith Press, 1988. 6. Sutherland DH, Kaufman KR, Moitoza JR. Kinematics of normal human walking. In:

Rose J, Gamble JG, eds. Human walking, 2nd ed. Baltimore: Williams & Wilkins, 1994;2:23-45.

7. Pathokinesiology Service, Physical Therapy Department. Normal and pathological gait syllabus. Downey, Calif.: Professional Staff Association of Rancho Los Amigos Hospital, 1977.

8. Ounpuu S, ed. Terminology for clinical gait analysis (Draft #2). Prepared by American Academy of Cerebral Palsy Developmental Medicine Gait Lab Committee and distributed at North American Clinical Gait Lab Conference, Benson Hotel, Portland, Ore., April 6-9, 1994.

9. Ayyappa E, ed. Words about words: the terminology of human walking, bipedal exchange. Monograph of the American Academy of Orthotists and Prosthetists Gait Society, Volumes 1-2, 1994.

10. Ayyappa E, ed. American Academy of Orthotists and Prosthetists Gait Society, Gait and Pathomechanics Syllabus, Certificate Program in Professional Development-Final Report, August 1996.

11. Finley FR, Cody K, Finizie R. Locomotive patterns in elderly women. Arch Phys Med Rehab 1969;50:140-6.

12. Ayyappa E. Gait lab technology: measuring the steps of progress. O&P Almanac 1996;45:2:28,29,41,42,56

13. Inman V. Ralston HJ, Todd F. Human walking. Baltimore: Williams & Wilkins, 1981 14. Bowker JH. Kinesiology and functional characteristics of the lower limb. In: Atlas of

limb prosthetics. St. Louis: CV Mosby, 1981;261-71 15. Rose J, Gamble J. Human walking, 2nd ed. Baltimore: Williams & Wilkins, 1994.

 

 

   

 

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Normal Human Locomotion, Part 2: Motion, Ground Reaction Force and Muscle Activity Edmond Ayyappa, CPO

ABSTRACT

In the patient care arena, an understanding of normal locomotion is a prerequisite to knowledge of pathological function in gait. Familiarity with joint motion, ground-reaction forces and muscular activity in normal individuals provides a bedrock of supporting knowledge that serves as a foundation for prosthetists and orthotists who seek to improve the performance of patients with pathological gait deficits. Although awareness of each of these components as it relates to a specific patient offers a revealing perspective by itself, in tandem they enable a three-dimensional differentiation between pathological and compensatory gait patterns. The sum is of far greater value than the individual parts.

In the past, successful orthotic and prosthetic intervention was limited primarily by design and material characteristics. A relatively gross understanding of gait mechanics was sufficient. With advancements in materials science and componentry development, such as the miniaturization of external power sources, the limitations to a patient's performance using a prosthesis or orthosis is more likely than ever to hinge on the practitioner's knowledge of gait mechanics.

This article, the second in a two-part series on normal human locomotion, attempts a narrative description of the dynamic phasic qualities of joint motion, ground-reaction forces and muscular activity.

Introduction

More than a century ago, the discipline of ambulatory performance assessment emerged as the components of human walking began to be measured and numerically quantified. The French scientist E.J. Marey performed pioneer research in gait assessment technology during the 1870s using multiple camera photography in sequenced series to assess movement, including pathological gait (see Figure 1) . His colossal contributions to gait science are revealed in his development of the first myograph for measuring muscle activity as well as a novel foot-switch measurement system for recording the magnitude and timing of plantar contact (1).

Eadweard Muybridge, an American contemporary and friend of Marey, was supported by Stanford University during the 1880s in using synchronized multiple camera photography with a scaled backdrop to capture movement on film and assess the motion of subjects walking (2). Other major advances into instrumented gait analysis were made by Scherb, who sought to understand the phasic action of muscle activity and performed hand-muscle palpation using a treadmill in 1920, and Adrian, who in 1925 advanced the use of EMG to study the dynamic action of muscles (3).

The applications of engineering and technology to the understanding of human walking received enormous impetus in 1945 when Inman et al. initiated the systematic collection of normal and amputee data on an instrumented walkway in their outdoor gait lab at the University of California--Berkeley (see Figure 2) (4-7). Since that time, a number of researchers and clinicians increasingly have used the growing array of gait technologies to measure and analyze the parameters of human performance in normal and pathological gait (8-12).

A contemporary instrumented walkway is a pathway that contains sensors or other measurement devices in the floor or around the subject's line of progression. The instrumentation is intended to monitor and measure one or more parameters of gait, such as motion (kinematics) or the forces creating motion (kinetics). Instrumentation placed directly on the patient also may measure phasic muscle activity (electromyographics or EMG), pressure against the skin or time-related (temporal) parameters such as velocity and cadence (13).

This article, presented as the second installment in the AAOP Professional Development Certificate Program in Gait and Pathomechanics (14,15), reviews the process of gait in phases from the perspective of kinematics, kinetics and EMG. To facilitate that goal, the basic principles of each are examined, and the current terminology (16) and means of data collection are reviewed.

Kinematics

Kinematics concerns itself with movement without consideration for the cause. The focus in gait analysis is on linear and angular displacements, velocities, accelerations and decelerations. The kinematics of walking can be quantitatively measured by means of instrumentation or qualitatively analyzed by means of observational gait assessment, such as a visual description of an individual's lower extremities, pelvis and trunk motion during ambulation. A qualitative contribution has been made by video technology using slow motion capabilities.

Recent development of inexpensive video gait assessment software packages that require manual measurements has clinical quantitative applications as well, enabling the measurement of joint angles in two dimensions (17). Since walking is a three-dimensional function, however, this type of video software may have limited value for comprehensive assessments or research applications that require a broad span of precise data. Methods that depend on hand measurements against a video image have questionable accuracy and are too labor-intensive for complex multijoint assessments. Also, the data collected from such methods cannot be integrated with kinetic data in real time or accurately displayed in real time or synchronized time with other data.

Most of what practitioners know today about kinematics of normal and pathological gait has been obtained from either an electrogoniometer, which uses electrical transducers attached to adjacent limb segments (see Figure 3) or, more likely, multicamera three-dimensional motion systems that track reflective markers placed on strategic anatomical landmarks.

Motion analysis measures dynamic range of motion. Dynamic range indicates joint motion or excursion from the maximum angle to the minimum angle during a particular phase or phases in the gait cycle. The computerized data obtained from multicamera motion systems can

document the motion of an individual's lower and upper extremities, pelvis, trunk, and head during ambulation. Motion analysis markers are small spheres or balls placed at specific bony landmarks that, when tracked by a camera-based video system, can be used to determine body segment and joint position. Active markers are joint and limb segment markers used during motion analysis that emit a light signal (see Figure 4) . Passive markers are markers that reflect visible or infrared light (see Figure 5) .

Kinetics

Kinetics is the general term given to the study of forces that cause movement. Force may be defined as a push or a pull and is produced when one object acts on another. The units used to measure force are Newtons (N). Forces in walking can be internal (such as muscle activity, ligamentous constraint, or friction in muscles and joints) or external (such as ground-reaction forces created from external loads).

Internal Moments

The rotational potential of the forces acting on a joint is called torque, moment or moment of force. The internal joint moment is the net result of all of the internal forces acting about the joint, including moments due to muscles, ligaments, joint friction and structural constraints. The joint moment usually is calculated around a joint center. When we think in terms of internal moments, for example, a net knee extensor moment means the knee extensors (quadriceps) are dominant at the knee joint, and the knee extensors are creating a greater moment than the knee flexors (hamstrings and gastrocnemius). The units used to express moments or torques are Newton-meters (N-m) and for research purposes usually are normalized to the subject's body mass. Normalization is the process by which a relationship is established between initially collected data (raw data) and some other basic reference data. Normalized to the subject's body mass, Newton-meters are expressed as N-m/kg.

The term joint power is used to describe the product of a joint moment and the joint angular velocity. Joint power is said to be generated when the moment and the angular velocity are in the same direction and said to be absorbed when they are in opposite directions. The units used to measure joint power are Watts (W).

Engineers and researchers studying gait attempt to measure the moment of force produced by muscles crossing a joint, the mechanical power flowing to and from those same muscles, and the energy changes of the body that result from this power. This requires the integration of both kinematic and kinetic data using very specialized data collection and processing systems (see Figure 6) .

External Moments

The external ground-reaction force line is a familiar concept to most clinicians trained in orthotics and prosthetics. Understanding its spatial relationship relative to the location of primary joints in normal gait is intuitively helpful in understanding the pathomechanics of a given patient.

A reaction force is the force that an initial body (A) exerts on a second body (B) in response to a force exerted by B on A. The reaction force has equal magnitude but opposite direction relative to the force exerted on A by B. Ground-reaction force is comprised of three

components: 1) vertical force, 2) fore-aft shear and 3) medial-lateral shear. Information on these forces is obtained from a force platform or force plate, which is a transducer set into the floor to measure the forces and torques applied by the foot to the ground (see Figure 7 and Figure 8 ).

These devices provide quantified measures of the three components of the resultant ground-reaction force vector and the resultant torque vector about a given joint. The ground-reaction force line essentially is the vector summation of the three reaction forces resulting from the interaction between the foot and ground. The moment of force or torque is the cross product of the radius vector and the force. The radius vector, traditionally assigned the variable r, is a position vector from the point around which the calculations are made to the line of action for the force being considered, traditionally assigned the variable F. The length of r is the moment arm of the force F. In two dimensions, the moment of force about a point is the product of a force and the perpendicular distance from the line of action of the force to the point. Typically, the moments of force are calculated about the center of rotation of a joint and are expressed in Newton-meters (N-m).

We have seen the resultant ground-reaction force (GRF) vector is the mean load bearing line, which takes into account both gravity and momentum (see Figure 9) . It has magnitude as well as directional qualities. The spatial relationship between this line and a given joint center influences the direction in which the joint will tend to rotate. This has enormous implications in understanding what orthotic or prosthetic component or alignment variant might be used to stabilize a joint during ambulation. The ground-reaction force line and the external moments or torques created at the major joints are presented in the pages that follow.

Electromyographics

Electromyographic (EMG) data provide important information in terms of understanding the direct physiological effect of prosthetic or orthotic design variants. Measuring muscle activity is like going straight to the mouth of the horse. Knowledge of the timing and intensity of the muscles throughout gait may suggest alterations in gait training and orthotic or prosthetic alignment or componentry to reduce excessive, ill-timed or prolonged muscle activity.

Electromyographic information is obtained by inserting fine wire electrodes directly into the muscle belly or by placing noninvasive surface electrodes over the muscle apex. Wire electrodes have the advantage of precise placement and are less likely to register "cross-talk" from adjacent muscles (see Figure 10) . Wire electrodes are essential for measuring deep muscles. Surface electrodes provide a noninvasive alternative for measuring muscle activity of superficial groups (see Figure 11) .

The Gait Cycle

Inman et al. (7) and Perry (8) have presented comprehensive models of human locomotion based on kinematic, kinetic and EMG collections, which have been drawn upon for the preparation of this article. The timing of the first five phases of gait identified by Perry can be precisely identified by the magnitude and timing of the vertical force graph of the stance limb. The typical free walk vertical force graph reveals itself as a curve with two peaks and a valley (see Figure 12) . By contrast, slow walk velocity and running do not display the midstance valley (F2) typical of normal locomotion in free walk velocity. The final three phases of gait that occur during swing can best be identified by kinematic positioning.

Initial Contact

The beginning of the gait cycle is referred to as initial contact. At the moment the foot strikes the ground, the ankle is at the neutral position, and the knee is close to full extension. In the sagittal plane, the alignment of the ground-reaction force vector at initial contact is posterior to the ankle joint, creating a plantarflexion moment (see Figure 13) . The three pretibial muscles (tibialis anterior, extensor digitorum longus and extensor hallicus longus), all of whose line of pull is anterior to the ankle joint, maintain the ankle and subtalar joint in neutral through eccentric contraction. The function of the peroneus tertius is considered identical to the extensor digitorum longus; they share the identical lateral tendon, and their muscle bellies blend into each other.

At the knee the vector is anterior to the joint axis, creating a passive extensor torque. Activity of the quadriceps and hamstring muscle groups continues from the previous terminal swing to preserve and stabilize the neutral position of the knee joint.

The hip and pelvis are emerging from a function of swing limb advancement with significant flexion, about 30 degrees. In normal gait, maximum hip flexion occurs during terminal swing and initial contact. A rapid high-intensity flexion moment thus is created at the hip as the vector falls anterior to the joint, placing great demand on the hip extensors. To restrain this impending flexion torque created by the anterior position of the vector, both the gluteus maximus and the hamstrings are activated. In the coronal plane, the gluteus medius is active preparing to stabilize the pelvis.

Loading Response

To absorb the impact force of loading and to maintain forward momentum, the eccentric action of the pretibial muscles regulates the ankle plantarflexion rate. A heel rocker action occurs as the pretibials pull the tibia forward over the fulcrum of the os calcis even as the foot is moving into a plantargrade position. This movement enables forward momentum of the tibia relative to the foot, but it also flexes the knee (see Figure 14) . During the peak of loading response, the magnitude of the vertical ground-reaction force exceeds body weight. The pretibials (tibialis anterior, extensor halicus longus and extensor digitorum longus) act as a shock absorber during loading response.

As a shock-absorbing mechanism and for energy efficiency, the knee flexes under the eccentric action of the quadriceps to about 15 to 18 degrees. During the stance phase of gait, the maximum knee-flexion angle usually is reached at foot flat. The quadriceps muscle group following this plantargrade posture controls the degree of knee flexion. Just as the pretibials advance the tibia forward over the foot in the rocker mechanism described, the quadriceps advance the femur over the tibia. This integrated action provides controlled forward movement of the entire lower-extremity unit.

The hip maintains its posture of about 30 degrees of flexion, creating a rapid, high-intensity flexion torque, the second-highest joint torque in normal gait after the dorsiflexion torque, which occurs at the talocrural joint during terminal stance. During loading response the hip extensors act as a shock absorber around the hip joint. Although not solely identified as hip extensors, the hamstrings nevertheless act as hip extensors as well as limit forward flexion of the pelvis and trunk. The function of the hamstrings when the hip is in flexion during stance is taken over by the gluteus maximus as stance progresses. Hip extensors prevent further flexion

at the hip, and shock absorption is provided by the gluteus maximus, hamstrings and adductor magnus. The medial-lateral control function of the hip adductors occurs as body weight is assumed by the stance leg.

Midstance

The momentum of forward progression over a stable foot with tibial stability maintained is referred to as the ankle rocker. The ankle rocker movement that progresses the tibia over a stationary foot is controlled early in midstance by the eccentric contraction of the soleus and is assisted by the gastrocnemius as the knee nears extension (see Figure 15) . At the beginning of midstance, the ankle is in a posture of 10 degrees of plantarflexion and moves through a range of more than 15 degrees to arrive at 5 to 7 degrees of dorsiflexion by the end of this phase. As the lower limb rolls forward over the stance foot, the body weight vector becomes anterior to the ankle joint, creating an increasing dorsiflexion moment. Activity of the soleus assisted by the gastrocnemius controls the rate of dorsiflexion. Action of the plantarflexors is crucial in providing limb stability as the contralateral toe-off transfers body weight onto the stance foot.

At the beginning of midstance, the vector is posterior to the knee joint but moves anterior as midstance progresses. The knee extends from 15 degrees of flexion to a neutral position. This is particularly mechanically efficient since plantarflexion of the ankle is most forceful with the knee in extension. The quads are active as knee extensors in early midstance only. Momentum of the contralateral swing leg creates an extension torque on the ipsilateral knee that decreases demand on the quadriceps and extends the knee without muscle action. By the end of midstance, the vector is anterior to the knee, creating passive stability. In the coronal plane, the ground-reaction force line is medial to the anatomical knee joint on the stance side, creating a varus moment. The moment is restrained by the capsular structures of the knee, especially the lateral collateral ligaments.

The hip joint is in a flexed posture of 30 degrees, which is reduced to 10 degrees as midstance progresses. The vector is anterior to the hip in early midstance and moves increasingly posterior to the hip, gradually reducing the flexion torque and diminishing the demand on the hip extensors. The gluteus maximus, active in early midstance, yields to this passive hip extension as the hip nears vertical alignment over the femur. Vertical ground-reaction force is reduced in magnitude at midstance due to the upward momentum of the contralateral swing limb. This upward momentum improves stability at the ipsilateral hip. The gluteus maximus, at this point not needed for sagittal stability, is active as an abductor rather than a hip extensor.

In the coronal plane, activity of hip abductors during midstance is essential to provide hip stability and avoid excessive pelvic tilt. In the frontal view the body mass and the ground-reaction force are quite medial to the structural support point at the head of the femur. At the time of midstance during gait it has been estimated that the vertical loading on the head of the femur on the stance side reaches a magnitude approximately equal to 21/2 times body weight (18). This creates a strong tendency toward excessive pelvic tilt (positive trendelenberg). The gluteus medius responds to limit pelvic tilt and stabilize the pelvis.

Terminal Stance

In terminal stance, forward fall of the body moves the vector further anterior to the ankle, creating a large dorsiflexion moment (see Figure 16) . Stability of the tibia on the ankle is provided by the eccentric action of the calf muscles. The plantarflexors are more active during this heel-off period than any other period of gait. The soleus and gastrocnemius prevent forward tibial collapse and allow the heel to rise over the metatarsal heads as the center of mass of the HAT (head, arms and trunk) advances over the foot. This is referred to as the forefoot rocker.

The forefoot rocker is comprised of two components, and some believe there are two distinct forefoot rockers. The initial forefoot rocker (third rocker) begins at heel off and ends when the contralateral limb contacts the ground. The mechanics are much different in the terminal forefoot rocker (fourth rocker), which occurs in preswing as body weight rapidly is unloading the ipsilateral limb and shifting to the contralateral side. The initial forefoot rocker (third rocker) serves as an axis around which progression of the body vector advances beyond the area of foot support, creating the highest demand of the entire gait cycle on the calf muscles. Minimal ankle movement of 5 degrees is required to reach 10 degrees of dorsiflexion, which then is maintained. The maximum amount of dorsiflexion of the anatomical ankle joint occurs during heel off.

The knee achieves an angular position of full extension accompanied by a mild extension torque that diminishes in the latter part of terminal stance. Joint stability and forward progression at the knee are achieved without muscle action.

While it once was believed the hip underwent up to 10 degrees of hyperextension during this period, it actually is likely to be less. The accuracy of early goniometric measurements at the hip is suspect. Electrogoniometers are not well-suited for measurement around the hip where they may be prone to reflect lumbar motion as well as soft tissue displacement. At any rate, hip extension combined with 5 degrees of pelvic rotation provides a smooth progression and facilitates an increased step length. A mild hip-extension torque is present. The trailing posture of the limb and the presence of the vector posterior to the hip provide passive stability at the hip joint. The tensor fascia lata serves to restrain the posterior vector at the hip. At the end of terminal stance, the magnitude of the vertical force reaches a second peak greater than body weight similar to that which occurred at the end of loading response.

Preswing

During preswing, the ankle moves rapidly from its dorsiflexion position at terminal stance to 20 degrees of plantarflexion (see Figure 17) . Although the ankle reaches its angular peak of plantarflexion during this period, actual plantarflexor activity is decreased in intensity as the limb is unloaded. In late preswing, the vertical force is diminished, and the plantarflexors are quiescent. There is no "push off" in normal reciprocal free walk bipedal gait. The dorsiflexion torque present at the beginning of preswing diminishes rapidly as the metatarsophalangeal joints extend to 60 degrees.

Passive knee flexion is created by planted hyperextended toes, advancement of the body past the metatarsal heads and contralateral loading. An early extension torque at the knee quickly gives way to a flexion torque. With the vector posterior to the knee, the knee flexes rapidly to achieve 35 degrees of flexion by the end of preswing, more than half the requirement for toe clearance in swing phase.

The hip flexes to a neutral position initiated by the rectus femoris, sartorius and adductor longus and assisted by momentum. The sagittal vector extends through the hip as the hip returns to a neutral posture. The adductor longus also decelerates the passive abduction created by contralateral body weight transfer. The continuing backward rotation of the pelvis effectively lengthens the trailing limb and counteracts hip flexion.

Initial Swing

Action of the pretibial muscles and long toe extensors begins to lift the foot and the ankle, which initially is at approximately 20 degrees of plantarflexion, its maximum achieved at any period in the gait cycle (see Figure 18) . By the end of initial swing, however, plantarflexion position is reduced to about 5 to 10 degrees, providing foot clearance for the midswing phase.

Although the knee began initial swing in only 30 degrees of flexion, the momentum from hip flexion assisted by the short head of the biceps femoris, sartorius and gracilis creates further rapid knee flexion to 60 degrees with the goal of providing limb advancement and foot clearance.

The hip is flexed 20 degrees initiated not only by the iliacus but by activity of both the gracilis and sartorius, which contribute to flexion of both the hip and knee joints.

Midswing

The knee extends as the ankle dorsiflexes, contributing to foot clearance while advancing the tibia (see Figure 19) . Pretibial muscle activity continues to preserve foot clearance as the ankle moves further toward dorsiflexion to reach a neutral position. Movement from plantarflexion toward dorsiflexion during the swing phase is referred to as dorsiflexion recovery.

Rapid knee extension, a passive event created by momentum, moves the knee from 60 to 30 degrees of flexion. Half of the knee extension needed for subsequent step length is achieved. The tibia assumes a relatively vertical position.

The hip flexors continue to preserve 30 degrees of hip flexion with mild EMG activity. The foot achieves ground clearance by 1 cm. The gracilis, sartorius and iliacus cut off in early midswing, and the hamstrings begin midway to decelerate the thigh. Additional limb advancement is created largely by momentum. Pelvic rotation is now neutral. The gluteus medius is quiescent on the ipsilateral side.

Terminal Swing

During terminal swing, the function of pretibial activity changes from one of foot clearance in swing to more appropriate limb placement and positioning for initial contact. A neutral position prepares the foot for the heel rocker function, assuring a heel first posture (see Figure 20) .

In the second half of terminal swing, the quadriceps extend the knee concentrically in a shortening contraction to facilitate full knee extension, which, assisted by pelvic rotation, accomplishes a full step length.

Eccentric contraction of both the hamstrings and the gluteus maximus is critical to accomplish deceleration of the thigh segment and restrain further hip flexion, which remains at 30 degrees. The long hamstrings have multiple roles of decelerating the leg, stabilizing the knee and limiting hip flexion in an eccentric or lengthening contraction. The gluteus maximus prepares for the impending forces of loading.

Conclusion

This review of human walking has explained several consistent patterns. With a few exceptions, muscular activity will oppose the external mechanical moment. Efficient body mechanics favors lengthening contractions, and agonists and antagonists active in opposition to each other actually are more the exception than the rule.

Normal human locomotion requires a complex interactive control between multiple limb and body segments that work congruently to provide the most shock-absorbing and energy-efficient forward movement possible. Gait characteristics are influenced by muscle strength, dynamic range of motion, and shape, position and function of numerous neuromuscular and musculoskeletal structures--as well as the ligamentous and capsular constraints of the joints. The primary goal is energy efficiency in progression using a stable kinetic chain of joints and limb segments that work congruently to transport the passenger unit forward.

Acknowledgements

The author would like to express appreciation to Ken Hudgens, program manager of the prosthetic and orthotic department, California State University--Dominguez Hills, for the preparation of Figure 9 .

 

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