Chapter 48, Factors Influencing Gait

CHAPTER

48 Characteristics of Normal Gait and Factors Influencing It THE GAIT CYCLE, THE BASIC UNIT OF GAIT . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .854 KINEMATICS OF LOCOMOTION . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .856 Temporal and Distance Parameters of a Stride . . . . . . . . . . . . . . . . . . . . . . . . . . . .856 Angular Displacements of Joints . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .857 MUSCLE ACTIVITY DURING LOCOMOTION . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .861 KINETICS OF LOCOMOTION . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .863 Joint Moments and Reaction Forces . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .863 Energetics of Gait: Power, Work, and Mechanical Energy . . . . . . . . . . . . . . . . . . .869 FACTORS THAT INFLUENCE PARAMETERS OF GAIT . . . . . . . . . . . . . . . . . . . . . . . . . . .872 Gender . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .872 Walking Speed . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .872 Age . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .872 SUMMARY . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .873

Habitual bipedal locomotion is a uniquely human function and influences an individual’s participation and interaction in society. Impairments in gait are frequent complaints of persons seeking rehabilitation services and are often the focus of an individual’s goals of treatment. Rehabilitation experts require a firm understanding of the basic mechanics of normal locomotion to determine the links between impairments of discrete segments of the musculoskeletal system and the patient’s abnormal movement patterns in gait. Therapists and other rehabilitation experts are called upon daily to analyze a patient’s movement and determine the cause of the abnormal, often painful, motion. A thorough understanding of normal locomotion and the factors that influence it, as well as an understanding of the functions of the components of the musculoskeletal system, provides a framework for evaluation and treatment of locomotor dysfunctions. This chapter describes the general characteristics of normal locomotion and introduces the clinician to the basic concepts central to all movement analysis. Normal human locomotion consists of stereotypical movement patterns that are immediately recognizable. Yet most individuals also are able to distinguish the gait of close friends and associates by the sound of their footsteps in the hallway. The purpose of this chapter is to describe the common characteristics of normal human locomotion and their variability and to provide insight into how impairments within the musculoskeletal

853

854

Part V | POSTURE AND GAIT

system may be manifested in altered gait patterns. The specific objectives of this chapter are to ■

Describe the basic components of the gait cycle

■

Present the temporal and distance characteristics of normal gait

■

Detail the angular displacement patterns of the joints of the lower extremity, the trunk, and the upper extremities

■

Describe the patterns of muscle activity that characterize normal locomotion

■

Briefly discuss the methods for determining muscle and joint loads sustained during normal locomotion and present the findings from representative literature

■

Briefly consider the energetics of normal locomotion and the implications of gait abnormalities on the efficiency of gait

Gait has been studied for millennia, and the last 50 years have seen an explosion in the research examining the characteristics of gait and the factors that control it. The current chapter is, of necessity, an overview of the characteristics of locomotion that are useful to a clinician and that demonstrate the effect of the integrity of the musculoskeletal system on gait. Several textbooks dealing only with locomotion provide details regarding the movement and methods of its assessment, and insight into the central nervous system’s role in controlling and modifying the movement of gait [28,119,127,159].

THE GAIT CYCLE, THE BASIC UNIT OF GAIT Gait is a cyclical movement that, once begun, possesses very repeatable events that continue repetitively until the individual begins to stop the motion. The steady-state movement of normal locomotion is composed of a basic repeating cycle, the gait cycle (Fig. 48.1). The cycle is traditionally defined as

Double support

Single support

Right step

the movement pattern beginning and ending with ground contact of the same foot. For example, using the right foot as the reference foot, the gait cycle begins when the right foot contacts the ground (usually with the heel) and ends when it contacts the ground again. Thus a gait cycle consists of the time the reference foot is on the ground (stance) and the time it is off the ground (swing). The movement of both limbs that occurs during the gait cycle is known as the stride.

Double support

Single support

Left step Stance

Swing Stride

Figure 48.1: The gait cycle of a single lower extremity consists of a stance and swing period and lasts from ground contact of one foot to the subsequent ground contact of the same foot. It includes two steps that are defined as the period from ground contact of one foot to the ground contact of the opposite foot. A single gait cycle includes two periods of double limb support and two periods of single limb support.

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

855

GC

FF

HO

CGC

TO

0%

15%

40%

50%

60%

Figure 48.2: The stance phase is divided into smaller phases that are demarcated by specific events. GC, ground contact; FF, foot flat; HO, heel off; CGC, contralateral ground contact, TO, toe off.

The stance phase of gait makes up approximately 60% of the gait cycle, so that the remaining 40% consists of the swing phase. The gait cycle with respect to the right limb is slightly out of phase with the gait cycle of the left limb. At contact on the right, the left limb is just ending its stance phase. At approximately 10% of the gait cycle on the right, the left limb leaves the ground and begins its swing phase, returning to the ground at approximately 50% of the gait cycle of the right limb. Thus the gait cycle is characterized by two brief periods, each lasting approximately 10% of the gait cycle, in which both limbs are in contact with the ground. These are periods of double limb support, and the remaining cycle consists of single limb support. The stance phase can be divided into smaller periods associated with specific functional demands and identified by distinct events (Fig. 48.2) [121]. The period immediately following ground contact is known as contact response, or weight acceptance, and ends when the whole foot flattens on the ground. During contact response, the limb absorbs the shock of impact and becomes fully loaded. The foot flat event that ends contact response occurs at approximately 15% of the normal gait cycle. It is important to recognize that loading response includes double limb support and continues into single limb support. The period following loading response is midstance, also known as trunk glide, since during this period the trunk glides over the fixed foot, moving from behind the stance foot to in front of it. Heel off ends trunk glide at approximately 40% of the gait cycle and begins terminal stance, which ends at 50% of the gait cycle when contralateral ground contact occurs. The final stage of stance, from 50 to 60% of the gait cycle, is preswing and is characterized by double limb support. It ends with toe off. The swing phase also is divided into early, middle, and late periods, although it lacks distinctive events to delineate these phases (Fig. 48.3). Early swing continues from 60% to approximately 75% of the gait cycle and is characterized by the rapid withdrawal of the limb from the ground. Midswing continues until approximately 85% of the gait cycle and consists of the period in which the swing limb passes the stance

60-75% Early swing

75-85% Mid swing

85-100% Late swing

Figure 48.3: The swing phase is divided into early swing, when the limb is pulled away from the ground; midswing, as the swing limb passes the stance limb; and late swing, when the swing limb extends toward the ground.

limb. Late, or terminal, swing finds the swing limb reaching toward the ground, preparing for contact. Although normal gait is often assumed to be symmetrical, substantial evidence exists to refute that assumption [13,61, 93,130]. Although the differences are small among ambulators without pathology, the right and left limb movements are not mirror images of one another. Differences exist in timing and movement patterns, in muscle activity, and in the loads applied to each limb [51,59]. When evaluating the gait patterns of individuals with asymmetrical impairments, clinicians must remember that small asymmetries in gait are normal. Consideration of the basic functional tasks of the swing and stance phase of gait provides a framework for characterizing the movements in each phase of gait. While the overriding goal of locomotion is forward progression, the stance and swing phases contribute to that goal in different ways. The stance phase has three tasks in locomotion: providing adequate support to avoid a fall, absorbing the shock of impact between the limb and the ground, and providing adequate forward and backward force for forward progress [35,158]. The basic tasks of the swing phase are safe limb clearance, appropriate limb

856

Part V | POSTURE AND GAIT

placement for the next contact, and transfer of momentum. By keeping these tasks in mind, the clinician can understand the importance of discrete movements of limb segments or the specific sequencing of muscle activity and can begin to appreciate the significance of specific joint impairments.

KINEMATICS OF LOCOMOTION As noted in Chapter 1, kinematics describes a movement in terms of displacement, velocity, and acceleration. The vast majority of kinematic analyses of gait examines displacement characteristics, and although velocity and acceleration data are available and may provide useful information, this chapter reviews the more commonly cited displacement data. Presented first is a description of the movement characteristics of the stride as a whole followed by descriptions of discrete movement patterns of individual joints. Many factors affect the kinematic characteristics of gait, including walking speed, age, height, weight or body mass index, strength and flexibility, pain, and aerobic conditioning. Walking speed and age have large and important effects on gait and are discussed later in this chapter. Unless noted otherwise, the data reported come from trials in which the subjects walk at their self-selected, comfortable, or free, speed.

Temporal and Distance Parameters of a Stride A stride consists of the movement of both limbs during a gait cycle and contains two steps. A step is operationally defined as the movement of a single limb from ground contact of one limb to ground contact of the opposite limb (Fig. 48.4). The literature demonstrates that there is considerable difference in step and stride characteristics among subjects and even among trials of the same subject [50]. Despite this normal variability, these parameters are capable of distinguishing between individuals with and without impairments [75,152].

TABLE 48.1

Step width Stride length

Foot angle Step length

Figure 48.4: Several distance measures help describe a typical gait cycle.

DISTANCE CHARACTERISTICS OF THE STRIDE The typical distance parameters of gait are defined in Table 48.1. A representative range of values also is presented from the literature [57,74,89,91,107,109,116]. Stride and step lengths depend directly upon standing height, so measures of absolute step or stride length, although frequently reported, are difficult to interpret. These measures can be normalized by standing height or lower extremity length to compare values from different individuals [27,73]. Estimates of normalized stride length vary from approximately 60 to 110% of standing height [27]. Judge et al. report a mean step length of 0.74
0.04 of leg length in young healthy adults [73]. Step width

Distance Parameters of Stride in Young Healthy Adults

Parameter

Definition

Range of Values Reported in the Literature

Stride length

The distance between ground contact of one foot and the subsequent ground contact of the same foot

1.33
0.09 to 1.63
0.11 m [57,74,89,91,107,109,116]

Step length

The distance between ground contact of one foot and the subsequent ground contact of the opposite foot

0.70
0.01 to 0.81
0.05 m [57,107,137]

Step width (also known as base of support)a

The perpendicular distance between similar points on both feet measured during two consecutive steps [25,104]

0.61
0.22 to 9.0
3.5 cm [91,107,109,137,146]

Foot angle

Angle between the long axis of the foot and the line of forward progression

5.1
5.7 to 6.8
5.6 [91]

a

Step width is defined variably in the literature. Some measures incorporate the angle of the foot on the ground.

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

TABLE 48.2

857

Temporal Parameters of Stride in Young, Healthy Adults

Parameter

Definition

Values from the Literature

Stride time

Time in seconds from ground contact of one foot to ground contact of the same foot

1.00
0.23 to 1.12
0.07 [91,107,109,116]

Speed (also known as velocity)

Distance/time, usually reported in m/sec

0.82–1.60
0.16 [47,57,73,74,89,91,106,109,116,137]

Cadence

Steps per minute

100–131 [27,40,47,73,74,91,106,109,137]

Stance time

Time in seconds that the reference foot is on the ground during a gait cycle

0.63
0.07 to 0.67
0.04 [91,107,109]

Swing time

Time in seconds that the reference foot is off the ground during a gait cycle

0.39
0.02 to 0.40
0.04 [91,107,109]

Swing/stance ratio

Ratio between the swing time and the stance time

0.63–0.64 [74,109]

Double support time

Time in seconds during the gait cycle that two feet are in contact with the ground

0.11
0.03 to 0.141
0.03 [91,106,107]

Single support time

Time in seconds during the gait cycle that one foot is in contact with the ground

Not reported

and foot angle are less frequently reported but provide an indication of the size of the base of support. TEMPORAL CHARACTERISTICS OF THE STRIDE The temporal characteristics of the stride are defined in Table 48.2 [40,47,53]. Included in this list is walking speed, or gait velocity, although this is typically computed over several strides. The normal gait cycle at free speed lasts approximately 1 second, and walking speed is between 3 and 4 miles per hour. Walking speed is a function of both cadence (steps/minute) and step length. An increase in either cadence or step length contributes to increased walking speed [6,54, 89,106,137,152]. Walking speed affects swing and stance time differently. Increased walking speed decreases the overall duration of the gait cycle, but the decrease in cycle duration results in a greater decrease in stance time than in swing time [6,106]. As stance time decreases with less change in swing time, double limb support time decreases, and single limb support time increases. The difference between running and walking is the absence of a double limb support phase in running. The ratio between swing and stance time increases toward 1 with increasing walking speed. Many gait disorders lead to altered time and distance parameters, typically decreased speed and stride length and, in the case of unilateral disorders, altered swing and stance times with abnormal swing–stance ratios. Such measures are relatively easy to obtain in the clinic and serve as useful outcome measures, sensitive to change. On the other hand, many different disorders produce similar temporal and distance characteristics. For example, a patient with unilateral hip pain and a patient with hemiparesis secondary to a stroke both walk with decreased velocity, and both demonstrate decreased single limb support time on the affected side and increased

double limb support time [106]. These parameters distinguish between normal gait and abnormal walking patterns but are unlikely to identify the differences in gait patterns between the two patients, even though such differences often are easily detected by an observer. Thus temporal and distance parameters may be helpful in tracking a patient’s progress but are insufficient to characterize a gait pattern fully and to identify the mechanisms driving the movement pattern. Patterns of joint excursions, however, can help the clinician to identify the differences in gait patterns between individuals with similar temporal and distance characteristics.

Angular Displacements of Joints The growth of photography in the mid- to late 19th century allowed the systematic observation of discrete movements of each joint during the complex activity of normal locomotion [5]. Over the last 50 years improved photographic techniques and the development of the computer have led to ever more precise monitoring of the three-dimensional motion of individual segments. The sagittal plane motions of the joints of the lower extremity are the most thoroughly studied and best understood, at least in part because sagittal plane motions are the largest and easiest to measure. In contrast, frontal and transverse plane motions of the joints of the lower extremities and the three-dimensional motions of the upper extremities and trunk are less frequently studied. Joint displacement data reveal intra- and intersubject variability in all planes, although the variability is greater in the frontal and transverse planes than in the sagittal plane and across subjects than between cycles of a single individual [14,39,55,74]. The smaller excursions in the frontal and transverse planes are particularly sensitive to differences in measurement procedure, which accounts for some of the increased variability of these motions [67]. Despite the variability in magnitudes of the movements, the patterns

858

Part V | POSTURE AND GAIT

and sequencing of joint movements in gait are remarkably consistent across trials and across subjects [12,31,32,106]. SAGITTAL PLANE MOTIONS

Rotation (degrees) Extension Flexion

The classic studies by Murray remain the foundation for understanding sagittal plane motion of the lower extremity [106,107,109,110] (Fig. 48.5). More-recent studies confirm

Ankle

Fl1

70

Fl2

90 110 Ex1

130 0

Ex2 20

Rotation (degrees) Extension Flexion

A

80

100

Knee

60 Fl2

40

Fl1

20 0

0

20

40 60 80 Percent of walking cycle

B

40

Ex2

Ex1

-20

Rotation (degrees) Extension Flexion

60 80 40 Percent of walking cycle

the overall patterns of motion for the hip, knee, and ankle, although there is variation in the reported maximal joint positions. Because studies demonstrate both intra- and intersubject variability, the reader is cautioned that the pattern of motion is the focus of the following discussion rather than the specific magnitudes [43,74,159]. Values of peak excursion are mentioned to provide an image of the motion rather than to define an absolute norm. The hip exhibits a single cycle of motion. Beginning at ground contact, the hip is in maximum flexion (approximately 25) and gradually extends, reaching maximum hip hyperextension (approximately 10) at close to 50% of the gait cycle, when contralateral ground contact occurs [80,87,106,159]. The magnitude of apparent hip hyperextension excursion depends on the point of reference. As noted in Chapter 38, a normal hip exhibits little or no hyperextension range of motion. Consequently, the hyperextension reported at the hip during locomotion is the result of pelvic motions in the transverse and sagittal planes. In most studies, the reported hip hyperextension reflects the orientation of the thigh with the trunk or with the roomfixed reference frame as seen in Fig. 48.6. After reaching maximum extension, the hip begins flexing again, reaching maximum flexion late in swing, at 80–85% of the gait cycle. The cycle repeats at ground contact. The knee exhibits a slightly more complex movement pattern, landing in extension, albeit usually a few degrees short of maximum extension, at ground contact. The knee flexes

100

Fl

Hip Fl1

60 80 100 120 Ex 0

C

25

50 75 Percent of walking cycle

100

Figure 48.5: Sagittal plane excursions of the ankle, knee, and hip (Reprinted with permission from Murray MP: Gait as a total pattern of movement. Am J Phys Med 1967; 46: 290–333.)

Figure 48.6: In most locomotion studies the hip excursion is described as the angle between the length of the thigh and a room-fixed coordinate system.

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

FRONTAL PLANE MOTIONS

Joint angle (degrees)

Frontal plane excursions are less well studied and more varied than sagittal plane movements (Fig. 48.7). Hip position in the frontal plane is affected by the motion of the pelvis over the femur and by the orientation of the femur as the subject translates toward the opposite foot to keep the center of mass over the base of support. The hip lies close to neutral abduction at ground contact and then adducts during weight acceptance as the pelvis drops on the contralateral side

40 Add 30 20 10 0 -10 Abd -20 20

A

Joint angle (degrees)

CLINICAL RELEVANCE: ASSOCIATED MOVEMENTS IN AN INDIVIDUAL FOLLOWING STROKE Close examination of the sagittal plane motions of the hip, knee, and ankle reveal that only for a very brief instant following toe off are these three joints moving in the same direction with respect to the ground. Just following toe off, all three joints are pulling the foot away from the ground, the hip and knee are flexing, and the ankle is dorsiflexing. At other points in the gait cycle the joints move independently, so that one or two joints move the foot toward the ground as the other(s) pull it away from the ground. A common impairment found in patients following stroke is an inability to disassociate movements, and as a result, a patient is compelled to move all three joints of the lower extremity together in the same direction. For example, to flex the knee, the patient may flex the knee and hip and dorsiflex the ankle simultaneously in a flexion pattern or extend the knee while simultaneously extending the hip and plantarflexing the ankle in an extension pattern. Such obligatory movements interfere with the normal timing and sequencing of joint movements in gait. For instance, in late swing, as the patient extends the knee toward the ground, the hip tends to extend, and the ankle plantarflex, producing a foreshortened step and an abnormal foot position at ground

contact. A flexion pattern produces similar conflicts as the hip begins to flex in terminal stance. At this time, the hip and knee should be flexing while the ankle continues to plantarflex. A flexion pattern stops the ankle plantarflexion and interferes with the normal roll off of late stance.

40 60 Gait cycle (%)

80

100

45 Add 30 15 0 Abd -15 20

B

40 Gait cycle (%)

60

Sup Subtalar rotation

10 to 20 immediately after contact, reaching maximum flexion at about 15% of the gait cycle when the subject achieves foot flat. At foot flat the knee begins to extend and reaches maximum extension at about 40% of the gait cycle as the heel rises from the ground. Flexion of the knee begins again and reaches a maximum of approximately 70 in midswing (approximately 75% of the gait cycle). Knee extension resumes, and the knee reaches maximum knee extension just before ground contact [20,80,88,106,129]. Ankle motion also exhibits several reversals in direction. Ground contact occurs with the ankle close to neutral in either slight plantarflexion or slight dorsiflexion [80,87,106]. Following contact, the ankle plantarflexes an additional 5 or 10, reaching a maximum at about 5% of the gait cycle. As the body glides over the stance foot, the ankle dorsiflexes, reaching a maximum just after the knee reaches full extension. Ankle plantarflexion resumes, and the ankle reaches maximum plantarflexion of approximately 20 just following toe off. In swing, the ankle dorsiflexes slightly but may remain in slight plantarflexion throughout swing. Pelvic motions in the sagittal plane are small, with no consistent definition of neutral. However, studies suggest that the pelvis anteriorly tilts whenever either hip is extending [107,109,145,147,154]. The anterior pelvic tilt contributes to the apparent hip hyperextension that occurs in late stance. Upper extremity sagittal plane motion also shows a rhythmic oscillation that is related to the movement of the lower extremities. At free walking speed, flexion of the shoulder and elbow parallel flexion of the opposite hip [106,110,151].

859

Toe off

80

100

Heel contact

10°

10° Pro 0

C

20

40 60 80 Percent of walking cycle

100

Figure 48.7: Frontal plane excursions of the hip (A), knee (B), and foot (C) are much smaller than sagittal plane excursions but show characteristic patterns of movement.

860

TRANSVERSE PLANE MOTIONS Transverse plane motions of the limbs and trunk also demonstrate more variability and smaller excursions than those seen in the sagittal plane (Fig. 48.10). Transverse plane rotations of the hip are a function of the transverse plane motion of the pelvis as well as the transverse plane motion of the femur (Fig. 48.11). Pelvic rotation in the transverse plane accompanies hip flexion, so that the pelvis rotates forward on the side of the flexing hip, reaching maximum forward rotation

Int

Hip Rotation

45 Joint angle (degrees)

[8,70,71,74,147] (Fig. 48.8). Adduction is amplified as the subject shifts toward the stance side to keep the center of mass over the foot. Adduction continues until late stance, when loading begins on the opposite limb. At that instance, the pelvis drops on the side in late stance, and the hip moves into abduction (Fig. 48.9). Reported knee motion in the frontal plane is slight, with estimates ranging from approximately 2 to 10 of adduction, peaking in early swing [8,20,74,88]. Frontal plane motion of the foot recorded during walking reflects the inversion and eversion component of supination and pronation of the foot. Although the position of the hindfoot at ground contact is variable and the magnitude of the reported excursions differs among reports, data consistently demonstrate a motion pattern characterized by eversion, consistent with pronation, following ground contact and continuing until mid to late stance when the hindfoot begins inverting or supinating [26,101,123,164]. Forefoot motion is similar to hindfoot motion, although forefoot pronation during stance begins after hindfoot pronation has begun [66,164].

Figure 48.9: During weight acceptance, the hip drops on the unsupported side, which is abducted.

CMC (w) = 0.882 CMC (b) = 0.845

30

15

0 20

A

40 60 Gait cycle (%)

80

100

80

100

0 Joint angle (degrees)

Figure 48.8: At weight acceptance, the individual shifts laterally to keep the center of mass close to the stance foot, and the pelvis drops on the unsupported side. The stance hip is in adduction.

Part V | POSTURE AND GAIT

Knee Rotation CMC (w) = 0.941 CMC (b) = 0.861

-10

-20 Ext -30 20

B

40 60 Gait cycle (%)

Figure 48.10: Transverse plane motions of the hip and knee. (Reprinted with permission from Kadaba MP, Ramakrishnan HK, Wootten ME, et al.: Repeatability of kinematic, kinetic, and electromyographic data in normal adult gait. J Orthop Res 1989; 7: 849–860.)

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

861

there is more disagreement about knee motion in swing [8,20,74,88,114,147]. Transverse plane motion of the knee is linked to the motion of the foot and to the sagittal plane motion of the knee, particularly during stance, when the lower extremity functions in a closed chain. As the foot pronates, the tibia medially rotates and allows the knee to flex. This coupled motion assists in shock absorption during loading response [122]. Later in stance, the foot supinates as the tibia rotates laterally, and the knee extends while the body rolls forward onto the opposite limb. MOTIONS OF THE TRUNK Studies of the head and trunk reveal that these segments undergo systematic translation and rotation in three dimensions and exhibit both intrasubject and intersubject variability [85,145,150]. The trunk exhibits slight flexion and extension during the gait cycle, is more erect or extended during single limb support, and is more flexed during double limb support [29,85]. Frontal plane motion of the trunk is consistent with the need to keep the center of mass over the stance foot. So the trunk leans slightly to the stance limb at each step [85,106,145,150]. In the transverse plane, the rotation of the trunk is opposite the rotation of the pelvis, with the trunk rotating forward on the side in which the shoulder is flexing [85,106,145]. Figure 48.11: The pelvic position in the transverse plane and the femoral rotation in the transverse plane both contribute to the transverse plane hip joint position during the gait cycle. At ground contact the femur is medially rotating, but the forward alignment of the pelvis contributes to lateral rotation of the hip. At heel off the opposite is true.

at approximately ground contact [74,106]. Forward rotation of the pelvis contributes to lateral rotation of the hip. At the same time, the opposite hip is in maximum extension, and the relative backward position of the pelvis on that side allows the hip to appear hyperextended. The transverse plane alignment of the pelvis on the extended hip tends to medially rotate the extended hip. Independent femoral movement provides its own contribution to hip position. At ground contact, the femur is aligned close to neutral but rotates medially from contact to midstance. Lateral femoral rotation then begins and continues into mid swing when medial rotation resumes. Hip joint position is the sum of the pelvic contribution and the femoral contribution to joint position. Although there is disagreement about the hip position at ground contact among the reported data, there is good consistency regarding the direction of the hip motion, medial rotation from ground contact to mid- or late stance and then lateral rotation until late swing or ground contact [70,71,74,114,147]. The knee, too, exhibits transverse plane motion with medial rotation following ground contact and gradual lateral rotation from midstance through most of swing, although

CLINICAL RELEVANCE: THE TRUNK’S CONTRIBUTION TO SMOOTH GAIT The gait pattern of a toddler learning to walk is characterized by large lateral leans with little forward rotation of the trunk and shoulders [9]. As the child matures, the pattern becomes smoother and more stable, and trunk rotation moves out of phase with the pelvis. The coupling motion of the trunk and pelvis contributes to the efficiency and stability of gait. Patients who lack the ability to rotate the trunk separately from the pelvis, such as patients with Parkinson’s syndrome or patients with low back pain, may lose gait efficiency and require more energy to walk.

MUSCLE ACTIVITY DURING LOCOMOTION Studies that examine the electrical activity of muscles during locomotion have played a central role in defining the role of muscles in producing and controlling locomotion. Data from Winter and Yack [163] demonstrate the normalized electromyographic (EMG) data for 16 muscles recorded in up to 19 subjects (Fig. 48.12). These data reveal important principles regarding muscle activity during gait. First, the duration of large bursts of activity for most muscles is quite brief, and most of these bursts occur at the transitions between swing and stance or between stance and swing. These data also

862

Part V | POSTURE AND GAIT

EMG normalized to mean X 100%

500 400 300 200 100 0

Gluteus medius N = 17 CV = 42% MEAN = 30.6

400 300

N = 16 CV = 58% MEAN = 16.5

Sartorius N = 15 CV = 54% MEAN = 25.5

200 100 0

600

Medial hamstrings N = 11 CV = 60% MEAN = 64.9

400 200

400 300 200

Rectus femoris N = 16 CV = 46% MEAN = 25.5

100 0

0 500 400 300 200 100 0

Lateral hamstrings N = 17 CV = 59% MEAN = 52.6

0

A 300

500 400 300 200 100 0

80 100

20 40 60 % of Stride

N = 11 CV = 51% MEAN = 33.9

Adductor longus

200

EMG normalized to mean X 100%

N = 11 CV = 38% MEAN = 37.0

Erector spinae

300

Gluteus maximus

200 100 0

400 300 200 100 0

300

0

0 400 300

Adductor magnus N = 11 CV = 55% MEAN = 42.7

0 Tibialis anterior

200

N = 12 CV = 28% MEAN = 135.4

N = 11 CV = 61% MEAN = 113.0

Lateral gastrocnemius

400 300

Soleus

200 100 0

100 0 300

Extensor digitorum longus N = 12 CV = 35% MEAN = 98.4

200

N = 10 CV = 57% MEAN = 79.2

80 100

N = 18 CV = 31% MEAN = 113

Peroneus longus N = 11 CV = 57% MEAN = 54.0

200 100 0

0 20 40 60 % of Stride

400 300

100

0

80 100

200 100 0

100

300

20 40 60 % of Stride

Medial gastrocnemius

200 100

200

N = 15 CV = 44% MEAN = 61.5

0

100

300

Vastus lateralis

0

20 40 60 % of Stride

80 100

B Figure 48.12: Electrical activity of lower extremity muscles during gait. (Reprinted with permission from Winter DA, Yack HJ: EMG profiles during normal human walking: stride-to-stride and inter-subject variability. Electroencephalogr Clin Neurophysiol 1987; 67: 402–411.)

demonstrate the considerable variability in muscle activity across individuals. Studies also demonstrate variability within a single individual, although there is less than across individuals [23,68,74,118,163]. Despite the variability of muscle activity, certain consistent functions for specific muscle groups emerge from the EMG data [10,41,68,74,81,103,138,155]. The gluteus maximus and hamstrings are active prior to and following ground contact, exerting a deceleration force on the hip and knee at the end of swing. Their activity also helps to initiate hip extension during early stance. The gluteus medius contracts just before ground contact and continues its activity through most of stance, until loading begins on the opposite side. The activity of the hip abductors provides essential frontal plane stability to the pelvis. The hip flexors contract in late stance and continue their activity into early swing to slow hip extension and initiate hip flexion. Muscle activity at the knee is characterized by cocontraction of the hamstrings and quadriceps for approximately the first 25% of the gait cycle, during loading response and early midstance. During this period, the knee is flexing and then extending, and the quadriceps activity is essential in controlling this movement. Some individuals exhibit activity of either the quadriceps, especially the rectus femoris, or hamstrings at the transition from stance to swing, but this activity is both variable and smaller in magnitude than the activity at the beginning of stance [7,112]. Most of swing proceeds with no muscle activity at the knee joint. The ankle also exhibits co-contraction of the dorsiflexor and plantarflexor muscles. Dorsiflexors of the ankle exhibit slight activity throughout swing to hold the foot away from the ground. The activity continues at ground contact and through the loading response, controlling the descent of the foot onto the ground. The plantarflexor muscles gradually increase their activity from ground contact through most of stance, with the greatest burst of activity from heel off to toe off as the body rolls over the plantarflexing foot. Review of the muscle activity of these large muscle groups demonstrates that much of the activity is characterized by an eccentric contraction followed by a concentric contraction. For example, the gluteus maximus contracts eccentrically as the hip flexes late in swing and then contracts concentrically as the hip begins to extend. The same pattern is found in the gluteus medius, hip flexors, quadriceps, and dorsiflexors. The plantarflexors also exhibit lengthening and then shortening, although at least some of the change in length is a passive stretch and shortening in the tendo calcaneus (Achilles tendon), so the actual change in muscle fiber length may be small [49]. The hamstrings also begin their activity with an eccentric contraction in late swing, but their subsequent length is more difficult to discern, since at loading response the hip is extending while the knee is flexing. The overall length change in the hamstrings during loading response may be negligible. The lengthening contractions that begin many

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

muscles’ activity in gait decelerate each joint, and then the subsequent concentric contractions begin the joint’s forward movement. It is worth noting that at most joints, the motion occurring during the concentric contraction continues after the contraction ceases. For example, the hip continues to extend long after the peak activity of the gluteus maximus and hamstrings, and the hip flexes after cessation of hip flexor activity. Similarly, the knee continues to extend without significant quadriceps activity, and the ankle continues to dorsiflex after the burst of dorsiflexor activity early in stance. Thus the chief functions of the muscles of the lower extremity during locomotion are to slow one motion and to provide an initial burst, or push, in the opposite direction. How motion continues in the absence of active muscle contraction is related to the kinetics of the movement.

863

undergo large linear accelerations, and joints exhibit large angular accelerations. As a result, the assumption used in static equilibrium analysis, that acceleration is negligible, is not valid when applied to gait. Newton’s second law of motion, 兺F  ma, states that the unbalanced force on a body is directly proportional to the acceleration of that body. The specific relationships between the accelerations and the forces and moments can be determined by applying the principles of dynamic equilibrium. The conditions of dynamic equilibrium are very similar to the conditions of static equilibrium. To determine the forces on an accelerating body in a two-dimensional analysis, the following conditions must be satisfied:

兺FX  maX, 兺FY  maY, 兺M  I  ␣

(Equation 48.1)

In three-dimensional analysis, the conditions for dynamic equilibrium are

KINETICS OF LOCOMOTION Kinetics examines the forces, moments, and power generated during a movement and, in the case of locomotion, includes the moments generated by the muscles, the forces applied across joints, and the mechanical power and energy generated. A discussion of the kinetics of gait allows consideration of the efficiency of gait.

Joint Moments and Reaction Forces As indicated in the preceding sections, gait consists of complex cyclical movements occurring in a coordinated sequence that is controlled by muscle activity. In addition, gait entails the repetitive impact loading of both lower extremities in each gait cycle. Thus it is easy to recognize that normal locomotion produces large forces between the foot and the ground, requires large muscle forces, and generates significant joint reaction forces. Many impairments in gait are related to an individual’s inability to generate sufficient muscular support or to sustain the large reaction forces of gait. DYNAMIC EQUILIBRIUM Researchers and clinicians have long been interested in the forces sustained by the muscles and joints during normal and abnormal locomotion [15,96,144]. Chapter 1 of this text describes the principles used to determine the loads in muscles and on joints during activity. Newton’s first law defines the conditions of static equilibrium (兺F  0, 兺M  0), stating that an object remains at rest (or in uniform motion) unless acted upon by an unbalanced external force. Throughout this text, two-dimensional examples of static equilibrium problems are provided to analyze the forces in the muscles and on joints during static tasks or in tasks where acceleration is negligible. However, during gait, limb segments

兺FX  maX, 兺FY  maY, 兺FZ  maZ

(Equation 48.2)

兺MX  I  ␣X, 兺MY  I  ␣Y, 兺MZ  I  ␣Z

(Equation 48.3)

and

where Fi is the force in the ith direction, a i is the linear acceleration in the ith direction, Mi is the moment about the ith axis, ␣i is the angular acceleration in the ith direction, and I is the moment of inertia. The moment of inertia indicates a body’s resistance to angular acceleration and depends on the body’s mass and distribution of mass. The larger the mass and the farther the mass is from the body’s center of mass, the larger is the body’s moment of inertia. Elite gymnasts tend to possess short and compact bodies (smaller moments of inertia) that allow high angular accelerations producing rapid rotations about horizontal bars and in tumbling routines. The acceleration quantities in each of the equations of dynamic equilibrium, mai and I  ␣I, are known as inertial forces and are intuitively explained by the awareness that it takes more force to push a car to start or stop its rolling than it takes to keep the car rolling. Solutions to the conditions of dynamic equilibrium, also known as equations of motion, require knowledge of several parameters, including mass and moment of inertia. Mass is usually determined from tables derived from cadaver measurements, as demonstrated in examples throughout this textbook [37]. Similarly, these tables provide means to calculate moments of inertia of a limb or limb segment from easily obtained anthropometric measurements, although methods also exist to compute the moment of inertia of some segments directly [22,134]. Regardless of the method chosen, the properties of mass and moment of inertia can be estimated and entered into the equations of motion to allow solutions.

864

Part V | POSTURE AND GAIT

Theoretically, the equations of motion in dynamic equilibrium can be used to calculate a body’s acceleration from all of the forces on the body. This approach is useful to determine the response of an airplane or rocket to an applied force. However, in the case of human movement, where forces cannot be measured directly, the equations of motion are used more often to determine the forces on the body when the accelerations are known. This approach, known as inverse dynamics, allows estimation of the forces on the human body and requires direct determination of the acceleration. Application of inverse dynamics in static equilibrium is straightforward because the accelerations are, by definition, zero, and the examples of two-dimensional analysis throughout this book demonstrate the use of inverse dynamics. Chapter 1 reminds the reader that acceleration is the change of velocity over time, and velocity is the change in displacement over time. Therefore, if a body’s displacement is known over time, then velocity and acceleration can be determined. Precise calculations of velocity and accelerations of the

body or of any limb segment requires careful measurement of the displacement, which can be accomplished by a number of techniques including high-speed cinematography, videography, or electromagnetic tracking devices [87,106, 111,124]. Appropriate signal processing of the displacement data and mathematical calculations yield satisfactory estimations of velocity and accelerations of the body of interest. A thorough discussion of the methods and challenges in these techniques is beyond the scope of this book; suffice it to say that the necessary acceleration values are available, so that the equations of motion can finally be solved for the applied forces. Examining the Forces Box 48.1 provides an example of the equations of motion for the leg–foot segment during the swing phase of gait. Using anthropometric data from Dempster [37], the mass (m) and moment of inertia (I) are entered directly into the calculations. Videographic data are collected at a rate of 60 Hz (hertz, or cycles per second) and manipulated so that the linear and angular accelerations of the leg–foot segment are determined for every 1/60 of a second

EXAMINING THE FORCES BOX 48.1 EQUATIONS OF MOTION IN TWO DIMENSIONS FOR THE LEG–FOOT SEGMENT DURING EARLY SWING

l3  the moment arm of the inertial force (maX) l4  the moment arm of the inertial force (maY) Since the limb segment accelerates during gait, the dynamic equilibrium conditions apply: FX  maX, FY  maY, M  I   where: aX, aY, and  are the x and y components of the linear accelerations and angular accelerations, respectively. These equations can be rewritten as FX  maX  0, FY  maY  0, M  I    0

may l4

J l2

l3

Fm

max l1 W

m  the mass of the leg and foot combined W  the weight of the leg and foot combined

where (maX), (maY), and (I   ) are known as inertial forces. The inertial forces contribute to moments about the knee joint so that taking moments about the knee, the motion equation is (W  l1)  (FM  l2)  [(maX)  l3]  [(maY)  l4] I Since the accelerations and anthropometric parameters, W and I, can be measured or determined from available data, the equation can be solved for the muscle force, F. Once the muscle force is determined, the joint reaction forces, JX and JY are calculated from:

FM  the muscle force

FX  maX

J  the joint reaction force

FMX  JX  maX

l1  the moment arm of the weight of the leg–foot

FY  maY

l2  the moment arm of the muscle

FMY  JY  W  maY

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

M

J

−Iα −ma

W

GRF Figure 48.13: Free body diagram of the leg–foot segment during stance includes the forces: weight of the leg-foot (W), joint reaction force (J), muscle force (M), ground reaction force (GRF), inertial forces ma and I, where m  mass, a  linear acceleration, I  moment of inertia, and   angular acceleration.

and entered into the equations. The equations of motion are solved repeatedly for the muscle force (F) at each increment of time. A similar procedure is applied to the stance phase of gait, but the external forces on the foot also include the ground reaction forces (Fig. 48.13). The direction and magnitude of these forces must be known to solve the equations of motion during stance and can be measured directly by force plates. The characteristics of the ground reaction force during gait are discussed in the following section. The example presented in Examining the Forces Box 48.1 assumes that only one muscle group is active. However, the EMG data described earlier in this chapter provide convincing evidence that there is co-contraction of the hamstrings and quadriceps during late swing and early stance and sometimes at the transition from late stance to early swing as well.

865

Thus there is more than one muscle applying force at the knee joint, producing a dynamically indeterminate system. As noted in Chapter 1 and elsewhere in this book, sophisticated mathematical solutions for indeterminate systems exist, and they are applied frequently in locomotion research to approximate the muscle and joint reaction forces [24,136]. Using inverse dynamics, many studies report the joint reaction forces in the body during the gait cycle [2,15,30, 42,60,83,136,141]. Peak joint reaction forces at the hip, knee, and ankle reported in the literature are presented in Table 48.3. These data reveal wide variation in the forces reported at each joint. Several factors influence these calculations, including the estimates of the body segment parameters of mass and moment of inertia, the accuracy of the displacement data and the procedures to determine accelerations, the use of two- or three-dimensional analysis, as well as the analytical approach used to complete the calculations [1,3,33,36, 84,166]. Values reported here are intended to demonstrate that regardless of the actual magnitude, all of the joints of the lower extremity sustain large and repetitive loads during locomotion. Running and jumping produce even larger muscle loads and joint reaction forces [18,98,149]. To avoid the problem of indeterminacy, researchers often solve only the moment equations, calculating the external moments applied to the limb by external forces such as weight and ground reaction forces and inferring the internal moments applied by the muscles and soft tissue [77]. Authors report either the internal [156] or external moment [80,86], and the reader is urged to read the literature carefully to identify which moment is reported. The limitation of this approach is that it prevents calculations of the forces in specific muscles and at the joints, but joint moments provide insight into the primary roles of muscle groups during gait and support the roles already suggested by EMG. Typical internal moments generated at the hip, knee, and ankle in the sagittal plane during normal locomotion are reported in Fig. 48.14. The internal moment at the hip joint at ground contact and contact response is an extension moment, consistent with the EMG activity of the gluteus maximus and hamstrings. The moment changes direction in midstance at about the time the hip extensors cease their activity and the flexors become active. The moment at the hip in swing is minimal until late swing when the hip extensors resume activity. The knee demonstrates a small and brief flexor moment at ground contact, consistent with hamstring activity, but then a larger and more prolonged extensor moment that is consistent

TABLE 48.3 Reported Peak Joint Reaction Forces during Normal Gait in Units of Body Weight Anderson et al. [2]

Komistek [83]

Duda et al. [42]

Seireg and Arvikar [136]

Hardt [60]

Simonsen et al. [141]

Hip

4

2.0–2.5

3

5.25

6

6

Knee

2.7

1.7–2.3

n.r.a

7

2.75

4.5

Ankle

6

1.25

n.r.

5

3.5

4

a

Not reported.

866

Part V | POSTURE AND GAIT

Hip Flexion/Extension Moment

% Nm/(BW•LL)

25 Flx

CMC (w) = 0.986 CMC (b) = 0.975

15 5 0 -5 Ext -15 20

A

40 60 Gait cycle (%)

80

100

Knee Flexion/Extension Moment % Nm/(BW•LL)

15 Flx CMC (w) = 0.960 CMC (b) = 0.944

5 0 -5 Ext

actually propel the body forward [120], recent studies provide convincing evidence that these muscles contribute some of the propulsion moving the body forward [21,113,125,132]. A very small dorsiflexion moment following toe off pulls the foot and toes away from the ground. Moments in the transverse and frontal planes also are reported and appear to be important in the mechanics and pathomechanics of locomotion [4,44,94]. However, less consensus exists regarding the magnitude and even the pattern of these moments. Moments in the frontal and transverse planes are smaller than those in the sagittal plane, and smaller moments are more sensitive to measurement errors, including the location of the joint axes and the kinematics of the movements [17,64]. Winter describes a support moment for the stance phase of gait that is the sum of the internal sagittal plane moments in which all of the moments that tend to push the body away from the ground or support the body are positive (Fig. 48.15) [63,156]. The net support moment during stance is positive, indicating the overall role of the muscles to support the body and to prevent collapse during weight bearing. Data suggest that although the net support moment is consistent across

-15 20

B

40 60 Gait cycle (%)

80

100

Ankle Flexion/Extension Moment

% Nm/(BW•LL)

10 Flx

CMC (w) = 0.992 CMC (b) = 0.981

0 MH -10 Ext -20 20

C

40 60 Gait cycle (%)

80

100

Figure 48.14: Internal moments at the hip, knee, and ankle in the sagittal plane. (Reprinted with permission from Kadaba MP, Ramakrishnan HK, Wootten ME, et al.: Repeatability of kinematic, kinetic, and EMG data in normal adult gait. J Orthop Res 1989; 7: 849–860.)

with quadriceps activity. In midstance, the knee exhibits a small flexor moment that is attributable to activity of the gastrocnemius. A small extension moment helps control knee flexion at the end of stance and in early swing, just as the flexion moment at the end of swing slows the rapid knee extension. A small dorsiflexion moment at ground contact and contact response reflects the dorsiflexor activity controlling the descent of the foot onto the ground. It is followed by a steadily increasing and prolonged plantarflexion moment controlling advancement of the tibia through the rest of stance. Although there has been disagreement about whether the plantarflexors

MK

MA

Figure 48.15: The support moment is the sum of the moments at the hip (MH ), knee (MK ), and ankle (MA ) needed to support the body weight during stance.

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

CLINICAL RELEVANCE: A PATIENT WITH QUADRICEPS WEAKNESS A patient with quadriceps weakness lacks the ability to support the knee actively during the stance phase of gait. To generate adequate support during the stance phase of gait, this patient may increase activity of the hip extensor muscles and of the soleus to increase the hip and ankle contributions to the net support moment (Fig. 48.16).

120 100 Force (% N/BW)

walking trials, individuals without pathology demonstrate variability in the individual joint moments, indicating that individuals with normal locomotor systems may exhibit flexibility in the ways they provide support [157].

867

Vertical Ant-Post Med-Lat

80 60 40 20 0 -20 20

40

60

80

100

Gait cycle (%)

GROUND REACTION FORCES With every stride, each foot applies a load to the ground and the ground pushes back, applying a ground reaction force to each foot. The magnitude and direction of this ground reaction force changes throughout the stance phase of each foot and is directly related to the acceleration of the body’s

Gluteus maximus

Soleus

Figure 48.16: An individual may increase the activity in the soleus and the gluteus maximus to support the knee in extension by preventing forward movement of the tibia or the femur, respectively.

Figure 48.17: Ground reaction forces during gait. (Reprinted with permission from Meglan D, Todd F: Kinetics of human locomotion. In: Rose J, Gamble JG, eds. Human Walking. Philadelphia: Williams & Wilkins, 1994; 23–44.)

center of mass. The center of mass of the body rises and falls as the individual moves from double support when the center of mass is low to single support when the center of mass is high [69,106,133]. Similarly, the center of mass moves from side to side as the individual passes from stance on the right to stance on the left [106]. The ground reaction force is measured directly by force plates imbedded in the walking surface. The ground reaction force typically is described by a vertical force as well as anterior–posterior and medial–lateral shear forces. The vertical ground reaction force under one foot is characterized by a double-humped curve (Fig. 48.17). The two peaks are greater than 100% of body weight and occur when the body accelerates upward. The valley between the peaks is less than 100% of body weight and occurs during single limb support. Examining the Forces Box 48.2 uses dynamic equilibrium to demonstrate how acceleration of the center of mass of the body alters the ground reaction force. The vertical ground reaction force also is characterized by a brief but high peak just following ground contact, which reflects the impact of loading [140]. CLINICAL RELEVANCE: GROUND REACTION FORCES AND JOINT PAIN Vertical ground reaction forces contribute significantly to joint reaction forces, and large joint reaction forces contribute to pain in patients with joint pathology such as arthritis. Patients with arthritis walk more slowly [76], and their vertical ground reaction forces demonstrate smaller peaks and valleys as the result of smaller vertical accelerations[139,142]. A reduction in walking speed, producing a reduction in accelerations, may be an effective way to reduce joint loads and, consequently, joint pain. These changes may represent appropriate adaptations to protect a painful joint and to maintain overall function.

868

Part V | POSTURE AND GAIT

EXAMINING THE FORCES BOX 48.2 THE CONTRIBUTION OF ACCELERATION TO THE VERTICAL GROUND REACTION FORCE Using the dynamic equilibrium condition, FY  maY, provides a direct demonstration of the role of the acceleration of the body’s center of mass in generating the vertical ground reaction force (GRF). a

FY  maY FY  maY  0 W  maY  GRF  0

a

GRF  W  maY When the body is accelerating toward the ground, the acceleration, aY, is negative, and the GRF is less than body weight, W. When the body accelerates upward away from the ground, acceleration, aY, is positive, and the GRF is greater than body weight, W.

The posterior and anterior shear components of the ground reaction force also demonstrate a consistent pattern in normal locomotion. The ground exerts a posterior force on the foot during the initial portion of stance, decelerating the foot; consequently this period is known as the deceleration phase. In midstance, the ground applies an anterior shear force on the foot, contributing to the forward propulsion of the body. The second half of the stance phase is known as the acceleration phase of the gait cycle. Walking on ice demonstrates the importance of these posterior and anterior shear forces. Because there is little friction between the foot and the ice, the posterior and anterior shear forces between the ground and the foot are small when walking on ice, and forward progress is impaired. In the absence of any posterior and anterior shear forces, forward progress is impossible. The medial and lateral shear forces during gait are smaller and more variable than the vertical forces or posterior–anterior shear forces. They reflect forces associated with the shift of the body from side to side between the supporting feet. Although plots of the ground reaction forces demonstrate rather stereotypical shapes, it is important to recognize that like kinematic variables, these forces exhibit normal intra- and intersubject variability [51,61]. The ground reaction force vector is the sum of the individual components of the ground reaction force. Whether described as a single force vector or as three individual components, the ground reaction force generates external moments on the joints of the body in all three planes (Fig. 48.18). Realistic computation of joint moments and forces during gait

GRF Figure 48.18: The ground reaction force vector (GRF) is the sum of the vertical, anterior–posterior, and medial–lateral ground reaction forces. The force vector applies external moments to the joints of the lower extremities about all three axes.

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

869

JOINT POWER

57% 55% 50% 45% 40% 35% 30% 25%

20% 15% 10% 5%

Mechanical power is the product of force and linear velocity or, in rotational motions such as the joint movements in locomotion, the product of joint moment and angular velocity: PMⴢ␻

where P is power in watts, M is a joint moment, and ␻ is the angular velocity of the limb segment. Power is a useful indication of the muscles’ role in controlling motion; it is negative when the body absorbs energy during eccentric muscle activity and is positive when the body generates energy during concentric muscle activity. Power also can be described as work (W) per unit time (t) (i.e., W兾t), where work is the product of force and displacement, or in angular terms, the product of moment (M) and angular displacement (␪): WMⴢ␪

2%

Figure 48.19: Progression of the center of pressure during locomotion. (Reprinted with permission from Sammarco GJ, Hockenbury RT: Biomechanics of the foot and ankle. In: Nordin M, Frankel VH, eds. Basic Biomechanics of the Musculoskeletal System. 3rd ed. Philadelphia: Lippincott Williams & Wilkins, 2001; 222–255.)

must include the three components of the ground reaction force or the force vector. The location of the ground reaction force with respect to the foot indicates the path of the center of pressure through the foot. In the normal foot, the center of pressure progresses in a relatively straight line from the posterior aspect of the plantar surface of the heel through the midfoot and onto the forefoot where it deviates medially onto the plantar surface of the great toe [56,58] (Fig. 48.19). Inability to roll over a painful toe or the interrupted forward progress of the body’s center of mass, because the knee suddenly hyperextends, are examples of gait deviations that produce changes in the pattern of the progression of the center of pressure.

Energetics of Gait: Power, Work, and Mechanical Energy Normal locomotion appears to be a remarkably efficient movement. Individuals without impairments, walking at a selfselected cadence, require less oxygen consumption than when walking at lower or higher cadences [65,104]. Individuals with locomotor impairments expend more energy during ambulation than individuals without impairments [16,97,143]. The efficiency of locomotion depends on many factors, including the mechanics of the muscular control of gait described earlier in this chapter and the conservation of mechanical energy that results from the synergistic movement of the limb segments.

(Equation 48.4)

(Equation 48.5)

Angular velocity, , is equal to angular displacement over time (  兾t) and therefore: P  M ⴢ 兾t

(Equation 48.6)

P  W兾t

(Equation 48.7)

and

Thus concentric muscle activity generates power, or does work, and eccentric activity absorbs power, and work is done on the segment [162]. A pogo stick (PogoTM) provides a useful example of positive and negative power, work done on or by the pogo stick (Fig. 48.20). In landing, the weight of the child does work on the pogo stick, and energy is absorbed by its spring, but in takeoff, the spring releases its energy and performs work on the child, pushing the child and pogo stick off the ground. Analysis of joint powers provides increasing understanding of the role of muscles in propelling and controlling movement during locomotion [21,125,130]. The joint powers at the hip, knee, and ankle during gait derived from two-dimensional analysis are pictured in Fig. 48.21. These demonstrate that positive power generation, when muscles are generating power and doing positive work, occurs at the hip at loading response as the hip extends and again at the end of stance as it flexes. Both of these periods are characterized by concentric muscle contractions. In contrast, the knee has only a brief period of power generation, producing only a small amount of power. Like the hip, the ankle generates considerable positive power at the end of stance when the plantarflexors contract concentrically. These data suggest that the hip flexors and extensors and the plantarflexors contribute important energy to the lower extremity during normal locomotion. A full understanding of the power generation and absorption in twoand three-dimensional analysis is still emerging and holds promise for providing more-direct insight into the mechanisms of gait deviations.

Part V | POSTURE AND GAIT

Power (% W/BW)

870 15 Gen 10

Hip

5 0 -5 Abs -10

Power (% W/BW)

0

20

40 60 Gait cycle (%)

80

15 Gen 10

100

Knee

5 0 -5 -10 -15 Abs -20

Power (% W/BW)

0

20

40 60 Gait cycle (%)

80

50 Gen 40

100

Ankle

30 20 10 0 Abs -10 -20

0

20

40 60 Gait cycle (%)

80

100

Figure 48.21: Joint powers at the hip, knee, and ankle from two-dimensional analysis. (Reprinted with permission from Meglan D, Todd F: Kinetics of human locomotion. In: Rose J, Gamble JG, eds. Human Walking. Philadelphia: Williams & Wilkins, 1994; 23–44.)

Figure 48.20: Energy storage and release. A. Weight bearing on the Pogo stick™ compresses its spring and work is done on the stick. B. As weight is removed, the spring is released, and the Pogo stick™ does work on the body, lifting it into the air.

CLINICAL RELEVANCE: JOINT POWERS IN INDIVIDUALS WITH GAIT DYSFUNCTIONS Joint powers during free-speed walking are altered in elders and in individuals with weaker lower extremity muscles [38,100]. The decrease in plantarflexion power and concomitant increase in hip flexor power generation noted in elders and in individuals with weakness may help to explain the decrease in velocity and step length reported in these individuals, as well as their mechanisms of compensation [38,99,100]. As an individual is unable to generate power through plantarflexion for forward progression, active hip flexion appears to provide the forward propulsion needed to swing the limb forward. These

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

patients may benefit from exercise to improve plantarflexion force production. The use of joint kinetics in conjunction with EMG is also useful in evaluating the complex gait deviations in individuals with central nervous system disorders such as cerebral palsy. These analyses provide more insight into the mechanics of the gait abnormalities than can be provided solely by clinical observation and lead to more- informed treatment decisions [117,128].

MECHANICAL ENERGY The cyclic movement inherent in locomotion and the ability of the muscles to store energy contribute to the inherent efficiency of normal gait. Mechanical energy, namely potential and kinetic energy, also provides insight into the efficiency of gait. Potential (PE) and kinetic (KE) energy are related to the distance of a body’s center of mass from the earth and to the body’s linear and angular velocity, as indicated by the following relationships: PE  mgh

(Equation 48.8)

where m is the mass of the body, g is the acceleration due to gravity, and h is the distance from the body’s center of mass to the earth; and 1 1 KE   mv 2   I 2 2 2

(Equation 48.9)

where m is the body’s mass, v is its linear velocity, I is its moment of inertia, and is its angular velocity. In an ideal system, conservation of energy dictates a complete transformation between potential and kinetic energy, so that an ideal roller coaster continues in motion indefinitely (Fig. 48.22). When the cars are at their peak height, potential energy is

Max PE Min KE

Min PE Max KE

Figure 48.22: In an ideal roller coaster, potential and kinetic energy are transformed from one form to the other with no loss of energy. Potential energy (PE  mgh) is maximum when the roller coaster is farthest from the ground, at the same time the kinetic energy (KE  1/2 mv2) is at its minimum. As the roller coaster descends the track it gains speed, increasing its kinetic energy while it is losing potential energy as it moves closer to the ground.

871

maximized and kinetic energy is minimized. At its lowest point, the roller coaster’s potential energy is minimum and its kinetic energy is maximum. Since the work done on a body equals the change in total energy, an ideal system requires no work to continue moving, since the change in the body’s total energy is zero. Studies of the mechanical energy of the limb segments during gait suggest that an exchange of kinetic and potential energy can account for most of the energy change in the distal leg at the beginning and end of swing [126,161]. This energy exchange improves when walking at free speed and is greater at steady-state walking than at the initiation of gait [95,102]. These studies demonstrate the efficiency of gait and how dependent the efficiency is on walking speed. The ability of the muscles to absorb and generate energy contributes to the overall efficiency of gait and explains how many of the movements can proceed without muscle contraction. Energy flows between adjacent limb segments during locomotion in much the same way that energy flows between the vaulter and the pole during a pole vault or among children playing “crack the whip.” Examination of the energy flow between limb segments reveals that the energy generated by the plantarflexors at push off is transferred passively to the leg and thigh, facilitating the initiation of swing. Similarly, the hamstrings absorb energy at the end of swing, and that energy is transferred to the trunk at ground contact, assisting in the trunk’s forward progression. The transfer of energy from segment to segment depends on the normal sequencing of the angular changes described earlier in this chapter. CLINICAL RELEVANCE: ENERGY TRANSFER AMONG LIMB SEGMENTS IN ABNORMAL GAIT Energy transfer among limb segments depends on the power generated and absorbed at joints and requires precise coordination among the moving segments. Since power is a function of the velocity of a limb segment, a limb segment that has a low angular velocity also has low power generation or absorption and, consequently, has less ability to transfer energy from one segment to another. A patient with arthritis producing a stiff knee is unable to transfer energy from the plantarflexors to the thigh; a patient with Parkinson’s disease, which is characterized by generalized rigidity, has difficulty transferring energy through the lower extremity and into the trunk because the joints lack the freedom of movement to allow the sequential movement patterns of the joints of the lower extremity. A study of patients with multiple sclerosis demonstrates an inverse relationship between the metabolic cost of walking and the patients’ ability to rapidly flex and extend the knee. This finding is consistent with a diminished capacity to transfer energy through the knee joint [115]. Thus treatments directed toward reducing joint stiffness or rigidity may lead to improved gait efficiency in these individuals.

872

FACTORS THAT INFLUENCE PARAMETERS OF GAIT Several factors influence gait performance and must be considered by clinicians evaluating and treating a person with a locomotor dysfunction. Factors considered here are gender, speed, and age.

Gender Although most observers would report differences between the gait patterns of males and females, few studies provide direct comparisons. Women walk with higher cadences than men and shorter strides [13,80,106]. Yet when the distance characteristics of the gait cycle are normalized by height, females demonstrate a similar or slightly larger stride length [45,80]. A study directly comparing 99 males and females of similar ages reports statistically different joint kinematics, although these differences are on the order of 2–4, and the clinical significance of these differences is negligible [80]. The same study also reports that females exhibit a statistically greater extension moment at the knee at initial contact and a greater flexion moment in preswing with increases in power absorption or generation at the hip, knee, and ankle. The authors suggest that these differences in kinetic measures may help to explain the higher incidence of knee osteoarthritis in women, but additional research is required to confirm these findings and demonstrate a clinical association.

Walking Speed Gait speed affects several parameters of gait performance. As noted in the discussion of the temporal and distance characteristics of gait, cadence, step length, and stride length increase with increased walking speed and decrease with decreased speed [6,106]. Increased speed appears to increase the variability of some temporal and spatial gait parameters such as step width [137]. Angular excursions also appear to increase with increased walking speed, although these changes are small and differ with the speed and joint examined [31,106,148]. Increases in joint excursions at the proximal joints are related to the increase in stride length associated with increased speeds [31]. Increased walking speeds also lead to increased ground reaction forces [6,25] and changes in the pattern of muscle activity. Although the relationship between muscle activity and walking speed is somewhat complex, there appears to be a general increase in the duration of muscle activity with increased walking speed, particularly in the muscles around the knee [103,105]. Similarly, joint moments and joint reaction forces increase with increased walking speed [11,135,166]. However, muscle activity during free-speed walking is more reproducible than that at speeds slower or faster than free speed [23,82]. Increased mechanical work and power at the knee and hip also accompany increased walking speed [21,72].

Part V | POSTURE AND GAIT

CLINICAL RELEVANCE: WALKING SPEED IN INDIVIDUALS WITH GAIT IMPAIRMENTS Many abnormal gait patterns found in individuals with impairments are characterized by decreased walking velocities. Patients with dysfunctions associated with low back pain, stroke, hemiparesis, and anterior cruciate ligament tears all frequently exhibit altered gait patterns that include decreased step length, smaller joint excursions, and decreased walking speed. Because decreased walking speed is associated with decreased step length and joint excursion, are the gait deviations exhibited by these patients merely the consequence of their walking speed? If a goal of treatment is to improve the gait pattern, the clinician must attempt to discern what characteristics of the gait pattern are attributable to the gait speed alone, and what characteristics are the result of the patient’s impairments.

Age Age appears to affect gait rather dramatically, as witnessed by the development of gait in the toddler and the apparent deterioration of gait in older adults. While the gradual acquisition of stable bipedal ambulation is a normal part of human development, it is unclear whether the alterations commonly seen in gait in the elderly are the normal consequence of aging or reflect the functional deficits resulting from impairments associated with neuromusculoskeletal disorders commonly found in elders [34,45,52,90,165]. Table 48.4 lists commonly reported changes in gait with aging. The ages of the elders studied range from approximately 60 years to over 100 years, and studies vary in the magnitude of changes reported. Despite the overwhelming data demonstrating changes in gait with increasing age, the nature of the relationship between age and locomotor function remains unclear. One of the most consistent findings with age is a decrease in free-walking speed [48,62,78,90, 92,108,153], but many of the other changes reported with aging also are consistent with the changes reported earlier in this chapter for walking speed alone [46]. Specifically, decreased walking speed produces reductions in step length, joint excursions, and ground reaction forces [48,62,79,160]. Consequently, many of the changes that occur with aging appear to be secondary changes associated with walking speed. Even the decrease in walking velocity reported with age appears to depend on an individual’s level of fitness and other factors besides age itself. Coexisting joint impairments; decreased strength of the quadriceps, plantarflexors, and hip flexors; hip and knee passive ranges of motion; and maximal oxygen uptake all help explain the diminished walking velocity seen with age [19,34,45,52,73,99]. Treatment of gait dysfunctions in elders requires consideration of the contributions made to the dysfunction by discrete impairments in the neuromusculoskeletal and cardiorespiratory systems.

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

TABLE 48.4

873

Commonly Reported Changes in Gait in Older Adults Change with Increased Age

Speed

Decreased [48,62,78,90,92,108,153]

Cadence

Increased [48,72]

Step/stride length

Decreased [46,48,62,72,73,108,160]

Double support time

Increased [46,72,160]

Joint angular excursions

Decreased [72,79,108] Unchanged [48]

Muscle activity

Increased [48]

Joint powers

Decreased generation in hip extension and plantarflexion and increased generation in hip flexion [72,73,79,160]

CLINICAL RELEVANCE: EVALUATION AND TREATMENT OF GAIT DYSFUNCTION IN ELDERS Data describing the gait of elderly individuals reveal that many of the changes thought to be characteristic of aging can be explained by a reduction in walking speed. Consequently, a clinician must alter the standards of “normal” used to judge the adequacy of gait. The gait patterns of elders walking at reduced speeds are not comparable to the patterns of subjects walking at faster speeds, regardless of age. Similarly, treatment may be most successful when directed toward those factors that contribute to diminished speed, including strength of the quadriceps, plantarflexors, and hip flexors.

SUMMARY This chapter reviews the kinematic and kinetic variables of normal gait. The kinematic variables presented in this chapter include the more global parameters of time and distance as well as the discrete displacement patterns of joints. Although all of these variables are subject to intra- and intersubject variability, representative values from the literature are presented to provide the reader with a frame of reference for normal locomotion. Joint excursions are largest in the sagittal plane and exhibit stereotypical patterns and sequences. In normal locomotion, the hip, knee, and ankle rarely move together toward or away from the ground. Activity of the major muscle groups of the lower extremity is reviewed. Their activity is typically brief, characterized by initial eccentric activity followed by concentric activity. In most cases, joint movement continues after muscle activity has ceased. The kinetic variables described in this chapter include ground and joint reaction forces, muscle forces, and joint moments, as well as joint power and mechanical energy. The principle of dynamic equilibrium is used to explain the derivation of muscle and joint reaction forces, joint moments, and joint power. Like the kinematic variables, the kinetic variables exhibit intra- and intersubject variability that reflects the normal variability of individuals and populations, but kinetic parameters also are quite sensitive to differences in meas-

urement procedures. The kinetic variables reveal that locomotion generates large muscle and joint forces. Kinetic analysis also demonstrates the remarkable efficiency of normal locomotion in which energy is stored and released, reducing the amount of work the muscles must perform to achieve the movement. Impairments in the neuromusculoskeletal system decrease the efficiency of gait. Finally this chapter discusses factors that influence walking patterns, including gender, walking speed, and age. The discussion reveals a complex interdependence between walking speed and age effects on gait, and the clinician is cautioned to keep these factors in mind when judging the walking performance of an individual.

References 1. Alkjaer T, Simonsen EB, Dyhre-Poulsen P: Comparison of inverse dynamics calculated by two- and three-dimensional models during walking. Gait Posture 2001; 13: 73–77. 2. Anderson FC, Pandy MG: Static and dynamic optimization solutions for gait are practically equivalent. J Biomech 2001; 34: 153–161. 3. Andrews JG: Methods for investigating the sensitivity of joint resultants to body segment parameter variations. J Biomech 1996; 29: 651–654. 4. Andriacchi TP: Dynamics of knee malalignment. Orthopedic Clinics of North America 1994; 25: 395–403. 5. Andriacchi TP, Alexander EJ: Studies of human locomotion: past, present and future. J Biomech 2000; 33: 1217–1224. 6. Andriacchi TP, Ogle JA, Galante JO: Walking speed as a basis for normal and abnormal gait measurements. J Biomech 1977; 10: 261–268. 7. Annaswamy TM, Giddings CJ, Della Croce U, Kerrigan DC: Rectus femoris: its role in normal gait. Arch Phys Med Rehabil 1999; 80: 930–934. 8. Apkarian J, Naumann S, Cairns B: A three dimensional kinematic and dynamic model of the lower limb. J Biomech 1989; 22: 143–155. 9. Assaiante C, Woollacott M, Amblard B: Development of postural adjustment during gait initiation: kinematic and EMG analysis. J Mot Behav 2000; 32: 211–226. 10. Battye CK, Joseph J: An investigation by telemetering of the activity of some muscles in walking. Med Biol Eng 1966; 4: 125.

874

11. Bergmann G, Deuretzbacher G, Heller M, et al.: Hip contact forces and gait patterns from routine activities. J Biomech 2001; 34: 859–871. 12. Bianchi L, Angelini D, Orani GP, Lacquaniti F: Kinematic coordination in human gait: relation to mechanical energy cost. Am Physiol Soc 1998; 79: 2155–2170. 13. Blanc Y, Balmer C, Landis T, Vingerhoets F: Temporal parameters and patterns of the foot roll over during walking: normative data for healthy adults. Gait Posture 1999; 10: 97–108. 14. Borghese NA: Kinematic determinants of human locomotion. J Physiol (Lond) 1996; 494: 863–879. 15. Bresler B, Frankel SP: The forces and moments in the leg during level walking. Trans ASME 1950; 27: 27–36. 16. Brown M, Hislop HJ, Waters RL, Porell D: Walking efficiency before and after total hip replacement. Phys Ther 1980; 60: 1259–1263. 17. Castagno P, Richards J, Miller F, Lennon N: Comparison of 3-dimensional lower extremity kinematics during walking gait using two different marker sets. Gait Posture 1995; 3: 87–87. 18. Cavanagh PR: The biomechanics of lower extremity action in distance running. Foot Ankle 1987; 7: 197–217. 19. Chang RW, Dunlop D, Gibbs J, Hughes S: The determinants of walking velocity in the elderly. Arthritis Rheum 1995; 38: 343–350. 20. Chao EY, Laughman RK, Schneider E, Stauffer RN: Normative data of knee joint motion and ground reaction forces in adult level walking. J Biomech 1983; 16: 219–233. 21. Chen IH, Kuo KN, Andriacchi TP: The influence of walking speed on mechanical joint power during gait. Gait Posture 1997; 6: 171–176. 22. Cheng CK, Chen HH, Chen CS, et al.: Segmental inertial properties of Chinese adults determined from magnetic resonance imaging. Clin Biomech 2000; 15: 559–566. 23. Chung SH, Giuliani CA: Within- and between-session consistency of electromyographic temporal patterns of walking in non-disabled older adults. Gait Posture 1997; 6: 110–118. 24. Collins JJ: The redundant nature of locomotor optimization laws. J Biomech 1995; 28: 251–267. 25. Cook TM, Farrell KP, Carey IA, et al.: Effects of restricted knee flexion and walking speed on the vertical ground reaction force during gait. J Orthop Sports Phys Ther 1997; 25: 236–244. 26. Cornwall MW, McPoil TG: Comparison of 2-dimensional and 3-dimensional rearfoot motion during walking. Clin Biomech 1995; 10: 36–40. 27. Craik RL, Dutterer L: Spatial and temporal characteristics of foot fall patterns. In: Craik RL, Oatis CA, eds. Gait Analysis: Theory and Application. St. Louis: Mosby-Year Book, 1995; 143–158. 28. Craik RL, Oatis CA: Gait Analysis: Theory and Application. St. Louis: Mosby, 1995. 29. Cromwell RL, Aadland TK, Nelson AT, et al.: Sagittal plane analysis of head, neck, and trunk kinematics and electromyographic activity during locomotion. J Orthop Sports Phys Ther 2001; 31: 255–262. 30. Cromwell RL, Schultz AB, Beck R, Warwick D: Loads on the lumbar trunk during level walking. J Orthop Res 1989; 7: 371–377. 31. Crosbie J, Vachalathiti R, Smith R: Age, gender and speed effects on spinal kinematics during walking. Gait Posture 1997; 5: 13–20.

Part V | POSTURE AND GAIT

32. Crosbie J, Vachalathiti R, Smith R: Patterns of spinal motion during walking. Gait Posture 1997; 5: 6–12. 33. Crowninshield RD, Brand RA: A physiologically based criterion of muscle force prediction in locomotion. J Biomech 1981; 14: 793–801. 34. Cunningham DA, Rechnitzer PA, Pearce ME, Donner AP: Determinants of self-selected walking pace across ages 19 to 66. J Gerontol 1982; 37: 560–564. 35. Das P, McCollum G: Invariant structure in locomotion. Neuroscience 1988; 25: 1023–1034. 36. Davy DT, Audu ML: A dynamic optimization technique for predicting muscle forces in the swing phase of gait. J Biomech 1987; 20: 187–201. 37. Dempster WT: Space requirements of the seated operator. In: Krogman WM, Fischer O, eds. Human Mechanics; Four Monographs Abridged AMRL-TDR-63-123. Wright-Patterson Air Force Base, OH: Behavioral Sciences Laboratory, 6570th Aerospace Medical Research Laboratories, Aerospace Medical Division, Air Force Systems Command, 1963; 215–340. 38. Devita P, Hortobagyi T: Age causes a redistribution of joint torques and powers during gait. J Appl Physiol 2000; 88: 1804–1811. 39. Dingwell JB, Cusumano JP: Nonlinear time series analysis of normal and pathological human walking. Am Inst Phys 2000; 10: 848–863. 40. Drillis R: Objective recording and biomechanics of pathological gait. Ann NY Acad Sci 1958; 74: 86–109. 41. Dubo HIC, Peat M, Winter DA, et al.: Electromyographic temporal analysis of gait: normal human locomotion. Arch Phys Med Rehabil 1976; 57: 415–420. 42. Duda GN: Internal forces and moments in the femur during walking. J Biomech 1997; 30: 933–941. 43. Dujardin FH, Roussignol X, Mejjad O, et al.: Interindividual variations of the hip joint motion in normal gait. Gait Posture 1997; 5: 246–250. 44. Eng JJ, Winter D, Patla AE: Intralimb dynamics simplify reactive control strategies during locomotion. J Biomech 1997; 30: 581–588. 45. Escalante A, Lichtenstein MJ, Hazuda HP: Walking velocity in aged persons: its association with lower extremity joint range of motion. Arthritis Care Res 2001; 45: 287–294. 46. Ferrandez AM, Paihous J, Durup M: Slowness in elderly gait. Exp Aging Res 1990; 16: 79–89. 47. Finley FR, Cody KA: Locomotion characteristics of urban pedestrians. Arch Phys Med Rehabil 1970; 51: 423–426. 48. Finley FR, Cody KA, Finizie RV: Locomotion patterns in elderly women. Arch Phys Med Rehabil 1969; 140–146. 49. Fukunaga T, Kubo K, Kawakami Y, et al.: In vivo behaviour of human muscle tendon during walking. Proc R Soc Lond B Biol Sci 2001; 268: 229–233. 50. Gabell A, Nayak USL: The effect of age on variability in gait. J Gerontol 1984; 39: 662–666. 51. Giakas G, Baltsopoulos V: Time and frequency domain analysis of ground reaction forces during walking: an investigation of variability and symmetry. Gait Posture 1997; 5: 189–197. 52. Gibbs J, Hughes S, Dunlop D, et al.: Predictors of change in walking velocity in older adults. J Am Geriatr Soc 1996; 44: 126–132. 53. Goodwin CS: The use of the voluntary muscle test in leprosy neuritis. Leprosy Rev 1968; 39: 209–216.

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

54. Grieve DW, Gear RJ: The relationships between length of stride, step frequency, time of swing and speed of walking for children and adults. Ergonomics 1966; 5: 379–399. 55. Growney E, Meglan D, Johnson M, et al.: Repeated measures of adult normal walking using a video tracking system. Gait Posture 1997; 6: 147–162. 56. Grundy M, Blackburn, Tosh PA, et al.: An investigation of the centres of pressure under the foot while walking. J Bone Joint Surg 1975; 57B: 98–103. 57. Hageman PA, Blanke DJ: Comparison of gait of young women and elderly women. Phys Ther 1986; 66: 1382–1386. 58. Han TR, Paik NJ, Im MS: Quantification of the path of center of pressure (COP) using an F-scan in-shoe transducer. Gait Posture 1999; 10: 248–254. 59. Hannah RE, Morrison JB, Chapman AE: Kinematic symmetry of the lower limbs. Arch Phys Med Rehabil 1984; 65: 65. 60. Hardt DE: Determining muscle forces in the leg during normal human walking—an application and evaluation of optimization methods. J Biomech Eng 1978; 100: 72–78. 61. Herzog W, Nigg BM, Read LJ, Olsson E: Asymmetries in ground reaction force patterns in normal human gait. Med Sci Sports Exerc 1989; 21: 110–114. 62. Himann JE, Cunningham DA, Rechnitzer PA, Paterson DH: Age-related changes in speed of walking. Med Sci Sports Exerc 1988; 20: 161–166. 63. Hof AL: On the interpretation of the support moment. Gait Posture 2000; 12: 196–199. 64. Holden JP, Stanhope S: The effect of variation in knee center location estimates on net knee joint moments. Gait Posture 1998; 7: 1–6. 65. Holt KG, Hamill J, Andres RO: Predicting the minimal energy costs of human walking. Med Sci Sports Exerc 1991; 23: 491–498. 66. Hunt AE, Smith RM: Inter-segment foot motion and ground reaction forces over the stance phase of walking. Clin Biomech 2001; 16: 592–600. 67. Hunt AE, Smith RM: Interpretation of ankle joint moments during the stance phase of walking: a comparison of two orthogonal axes systems. J Appl Biomech 2002; 17: 173–180. 68. Hunt AE, Smith RM, Torode M: Extrinsic muscle activity, foot motion and ankle joint moments during the stance phase of walking. Foot Ankle Int 2001; 22: 31–41. 69. Iida H, Yamamuro T: Kinetic analysis of the center of gravity of the human body in normal and pathological gait. J Biomech 1987; 20: 987–995. 70. Isacson J, Gransberg L, Knutsson E: Three-dimensional electrogoniometric gait recording. J Biomech 1986; 19: 627–635. 71. Johnston RC, Smidt GL: Measurement of hip joint motion during walking: evaluation of an electrogoniometric method. J Bone Joint Surg 1969; 51A: 1083. 72. Judge J, Davis RB, Ounpuu S: Age-associated reduction in step length: testing the importance of hip and ankle kinetics. Gait Posture 1995; 3: 81–81. 73. Judge JO: Step length reductions in advanced age: the role of ankle and hip kinetics. J Gerontol A Biol Sci Med Sci 1996; 51: M303–M312. 74. Kadaba MP, Ramakrishnan HK, Wootten ME, et al.: Repeatability of kinematic, kinetic, and electromyographic data in normal adult gait. J Orthop Res 1989; 7: 849–860.

875

75. Katoh Y, Chao YS, Laughman RK, et al.: Biomechanical analysis of foot function during gait and clinical applications. Clin Orthop 1983; 177: 23–33. 76. Kaufman K, Hughes C, Morrey BF, et al.: Gait characteristics of patients with knee osteoarthritis. J Biomech 2001; 34: 907–915. 77. Kepple TM, Siegel KL, Stanhope SJ: Relative contributions of the lower extremity joint moments to forward progression and support during gait. Gait Posture 1997; 6: 1–8. 78. Kernozek TW, LaMott EE: Comparisons of plantar pressures between the elderly and young adults. Gait Posture 1995; 3: 143–148. 79. Kerrigan DC: Biomechanical gait alterations independent of speed in the healthy elderly: evidence for specific limiting impairments. Arch Phys Med Rehabil 1998; 79: 317–322. 80. Kerrigan DC: Gender differences in joint biomechanics during walking: normative study in young adults. Am J Phys Med Rehabil 1998; 77: 2–7. 81. Kleissen RFM, Litjens MCA, Baten CTM, et al.: Consistency of surface EMG patterns obtained during gait from three laboratories using standardised measurement technique. Gait Posture 1997; 6: 200–209. 82. Knutson LM, Soderberg GL: EMG: use and interpretation in gait. In: Craik RL, Oatis CA, eds. Gait Analysis: Theory and Application. St. Louis: Mosby, 1995; 307–325. 83. Komistek RD, Stiehl JB, Dennis DA, et al.: Mathematical model of the lower extremity joint reaction forces using Kane’s method of dynamics. J Biomech 1998; 31: 185–189. 84. Koopman B, Grootenboer HJ, de Jongh HJ: An inverse dynamics model for the analysis, reconstruction and prediction of bipedal walking. J Biomech 1995; 28: 1369–1376. 85. Krebs DE, Wong D, Jevsevar D, et al.: Trunk kinematics during locomotor activities. Phys Ther 1992; 72: 505–514. 86. Kuruvilla A, Costa JL, Wright RB, et al.: Characterization of gait parameters in patients with Charcot-Marie-Tooth disease. Neurol India 2000; 48: 49–55. 87. Kuster M, Sakurai S, Wood GA: Kinematic and kinetic comparison of downhill and level walking. Clin Biomech 1995; 10: 79–84. 88. Lafortune MA, Cavanagh PR, Sommer HJ, Kalenak A: Threedimensional kinematics of the human knee during walking. J Biomech 1992; 25: 347–357. 89. Larsson L-E, Odenrick P, Sandlund B, et al.: The phases of the stride and their interaction in human gait. Scand J Rehabil Med 1980; 12: 107–112. 90. Leiper CI, Craik RL: Relationships between physical activity and temporal-distance characteristics of walking in elderly women. Phys Ther 1991; 71: 791–803. 91. Lemke MR, Wendorff T, Mieth B, et al.: Spatiotemporal gait patterns during over ground locomotion in major depression compared with healthy controls. J Psychiatr Res 2000; 34: 277–283. 92. Lundgren-Lindquist B, Aniansson A, Rundgren A: Functional studies in 79-year olds: III. Walking performance and climbing capacity. Scand J Rehabil Med 1983; 15: 125–131. 93. Macellari V, Giacomozzi C, Saggini R: Spatial-temporal parameters of gait: reference data and a statistical method for normality assessment. Gait Posture 1999; 10: 171–181. 94. MacKinnon CD, Winter DA: Control of whole body balance in the frontal plane during human walking. J Biomech 1993; 26: 633–644.

876

95. Mansour JM, Lesh MD, Nowak MD, Simon SR: A three dimensional multi-segmental analysis of the energetics of normal and pathological human gait. J Biomech 1982; 15: 51–59. 96. Mansour JM, Pereira JM: Quantitative functional anatomy of the lower limb with application to human gait. J Biomech 1987; 20:1: 51–58. 97. Martin PE, Morgan DW: Biomechanical considerations for economical walking and running. Med Sci Sports Exerc 1992; 24: 467–474. 98. McClay I, Manal K: Three-dimensional kinetic analysis of running: significance of secondary planes of motion. Med Sci Sports Exerc 1999; 31: 1629–1637. 99. McGibbon CA, Krebs DE: Effects of age and functional limitation on leg joint power and work during stance phase of gait. J Rehabil Res Dev 1999; 36: 173–182. 100. McGibbon CA, Puniello MS, Krebs D: Mechanical energy transfer during gait in relation to strength impairment and pathology in elderly women. Clin Biomech 2001; 16: 324–333. 101. McPoil T, Cornwall MW: Relationship between neutral subtalar joint position and pattern of rearfoot motion during walking. Foot Ankle 1994; 15: 141–145. 102. Miller CA, Verstraete MC: A mechanical energy analysis of gait initiation. Gait Posture 1999; 9: 158–166. 103. Milner M, Basmajian JV, Quanbury AO: Multifactorial analysis of walking by electromyography and computer. Am J Phys Med 1971; 50: 235 104. Minetti AE, Capelli C, Zamparo P, et al.: Effects of stride frequency on mechanical power and energy expenditure of walking. Med Sci Sports Exerc 1995; 27: 1194–1202. 105. Miyashita M, Matsui H, Miura M: The relation between electrical activity in muscle and speed of walking and running. Med Sport 1971; 6: 192–196. 106. Murray MP: Gait as a total pattern of movement. Am J Phys Med 1967; 46: 290–333. 107. Murray MP, Drought AB, Kory RC: Walking patterns of normal men. J Bone J Surg 1964; 46: 335. 108. Murray MP, Kory RC, Clarkson BH: Walking patterns in healthy old men. J Gerontol 1969; 24: 169. 109. Murray MP, Kory RC, Sepic SB: Walking patterns of normal women. Arch Phys Med 1979; 51: 637. 110. Murray MP, Sepic SB, Barnard EJ: Patterns of sagittal rotation of the upper limbs in walking. Phys Ther 1967; 47: 272–284. 111. Nawoczenski DA, Baumhauer JF, Umberger BR: Relationship between clinical measurements and motion of the first metatarsophalangeal joint during gait. J Bone Joint Surg Am 1999; 81: 370–376. 112. Nene A, Mayagoitia R, Veltink P: Assessment of rectus femoris function during initial swing phase. Gait Posture 1999; 9: 1–9. 113. Neptune RR, Kautz SA, Zajac FE: Contributions of the individual ankle plantarflexors to support, forward progression and swing initiation during walking. J Biomech 2001; 34: 1387–1398. 114. Nester C: The relationship between transverse plane leg rotation and transverse plane motion at the knee and hip during normal walking. Gait Posture 2000; 12: 251–256. 115. Olgiatti R, Burgunder JM, Mumenthaler M: Increased energy cost of walking in multiple sclerosis: effect of spasticity, ataxia, and weakness. Arch Phys Med Rehabil 1988; 69: 846–849.

Part V | POSTURE AND GAIT

116. Ostrosky KM, VanSwearingen JM, Burdett RG, Gee Z: A comparison of gait characteristics in young and old subjects. Phys Ther 1994; 74: 637–646. 117. Ounpuu S, Thompson JD, Davis RB III, DeLuca PA: An examination of the knee function during gait in children with myelomeningocele. J Pediatr Orthop 2000; 20: 629–633. 118. Ounpuu S, Winter D: Bilateral electromyographical analysis of the lower limbs during walking in normal adults. Electroencephalogr Clin Neurophysiol 1989; 72: 429–438. 119. Patla AE: Adaptability of Human Gait—Implications for the Control of Locomotion. New York: Elsevier /North-Holland, 1991. 120. Perry J: Ankle foot complex. In: Gait Analysis: Normal and Pathological Function. Thorofare, NJ: Slack, 1992; 51–87. 121. Perry J: Hip. In: Gait Analysis, Normal and Pathological Function. Thorofare, NJ: Slack Incorporated, 1992; 119. 122. Perry SD, Lafortune MA: Influences of inversion/eversion of the foot upon impact loading during locomotion. Clin Biomech 1995; 10: 253–257. 123. Pierrynowski MR, Smith SB: Rear foot inversion/eversion during gait relative to the subtalar joint neutral position. Foot Ankle Int 1996; 17: 406–412. 124. Reinschmidt C, van den Bogert AJ, Lundberg A, et al.: Tibiofemoral and tibiocalcaneal motion during walking: external vs. skeletal markers. Gait Posture 1997; 6: 98–109. 125. Riley PO, Dellacroce U, Kerrigan DC: Effect of age on lower extremity joint moment characteristics to gait speed. Gait Posture 2001; 14: 264–270. 126. Robertson DGE, Winter DA: Mechanical energy generation, absorption, and transfer amongst segments during walking. J Biomech 1980; 13: 845–854. 127. Rose J, Gamble JG: Human Walking. Baltimore: Williams & Wilkins, 1994. 128. Rose SA, DeLuca PA, Davis RB III, Ounpuu S: Kinematic and kinetic evaluation of the ankle after lengthening of the gastrocnemius fascia in children with cerebral palsy. J Pediatr Orthop 1993; 13: 727–732. 129. Rowe PJ, Myles CM, Walker C, Nutton R: Knee joint kinematics in gait and other functional activities measured using flexible electrogoniometry: how much knee motion is sufficient for normal daily life? Gait Posture 2000; 12: 143–155. 130. Sadeghi H, Allard P, Duhaime M: Contributions of lower-limb muscle power in gait of people without impairments. Phys Ther 2000; 80: 1188–1196. 131. Sadeghi H, Allard P, Prince F, Labelle H: Symmetry and limb dominance in able-bodied gait: a review. Gait Posture 2000; 12: 34–45. 132. Sadeghi H, Sadeghi S, Prince F, et al.: Functional roles of ankle and hip sagittal muscle moments in able-bodied gait. Clin Biomech 2001; 16: 688–695. 133. Saini M, Kerrigan DC, Thirunarayan MA, Duff-Raffaele M: The vertical displacement of the center of mass during walking: a comparison of four measurement methods. J Biomech Eng 1998; 120: 133–139. 134. Sarfaty O, Ladin Z: A video-based system for the estimation of the inertial properties of body segments. J Biomech 1993; 26: 1011–1016. 135. Scott SH, Winter DA: Internal forces at chronic running injury sites. Med Sci Sports Exerc 1990; 22: 357–369.

Chapter 48 | CHARACTERISTICS OF NORMAL GAIT AND FACTORS INFLUENCING IT

136. Seireg A, Arvikar RJ: The prediction of muscular load sharing and joint forces in the lower extremities during walking. J Biomech 1975; 8: 89–102. 137. Sekiya N, Nagasaki H, Ito H, Furuna T: Optimal walking in terms of variability in step length. J Orthop Sports Phys Ther 1997; 26: 266–272. 138. Shiavi R: Electromyographic patterns in adult locomotion: a comprehensive review. J Rehabil Res Dev 1985; 22: 85–98. 139. Simkin A: The dynamic vertical force distribution during level walking under normal and rheumatic feet. Rheumatol Rehabil 1981; 20: 88–97. 140. Simon SR, Paul IL, Mansour J, et al.: Peak dynamic force in human gait. J Biomech 1981; 14: 817–822. 141. Simonsen EB, Dyhre-Poulsen P, Voigt M, et al.: Bone-on-bone forces during loaded and unloaded walking. Acta Anat (Basel) 1995; 152: 133–142. 142. Smidt GL, Wadsworth JB: Floor reaction forces during gait: comparison of patients with hip disease and normal subjects. Phys Ther 1973; 53: 1056. 143. Song KM: The effect of limb-length discrepancy on gait. J Bone Joint Surg Am 1997; 79: 1690–1698. 144. Stauffer RN, Chao EYS, Brewster RC: Force and motion analysis of the normal, diseased and prosthetic ankle joint. Clin Orthop 1977; 127: 189–196. 145. Stokes VP, Andersson C, Frossberg H: Rotational and translational movement features of the pelvis and thorax during adult human locomotion. J Biomech 1989; 22: 43–50. 146. Stolze H, Kuhtz-Buschbeck JP, Mondwurf C, et al.: Retest reliability of spatiotemporal gait parameters in children and adults. Gait Posture 1998; 7: 125–130. 147. Sutherland DH, Kaufman KR, Moitoza JR: Kinematics of normal human walking. In: Rose J, Gamble JG, eds. Human Walking. Philadelphia: Williams & Wilkins, 1994; 23–44. 148. Taylor NF, Goldie PA, Evans OM: Angular movements of the pelvis and lumbar spine during self-selected and slow walking speeds. Gait Posture 1999; 9: 88–94. 149. van den Bogert AJ, Read L, Nigg BM: An analysis of hip joint loading during walking, running, and skiing. Med Sci Sports Exerc 1999; 31: 131–142. 150. Vogt L, Banzer W: Measurement of lumbar spine kinematics in incline treadmill walking. Gait Posture 1999; 9: 18–23. 151. Wagenaar RC, van Emmerick REA: Resonant frequencies of arms and legs identify different walking patterns. J Biomech 2000; 33: 853–861.

877

152. Wall JC, Devlin J, Khirchof R, Lackey B: Measurement of step widths and step lengths: a comparison of measurements made directly from a grid with those made from a video recording. J Orthop Sports Phys Ther 2000; 30: 410–417. 153. Waters RL, Hislop HJ, Perry J, et al.: Comparative cost of walking in young and old adults. J Orthop Res 1983; 1: 73–76. 154. Whittle MW, Levine DF: Sagittal plane motion of the lumbar spine during normal gait. Gait Posture 1995; 3: 82. 155. Winter D: Biomechanics and Motor Control of Human Movement. New York: John Wiley & Sons, 1990. 156. Winter DA: Overall principle of lower limb support during stance phase of gait. J Biomech 1980; 13: 923–927. 157. Winter DA: Kinematic and kinetic patterns in human gait: variability and compensating effects. Hum Move Sci 1984; 3: 51–76. 158. Winter DA: Biomechanics of normal and pathological gait: implications for understanding human locomotor control. J Mot Behav 1989; 21: 337–355. 159. Winter DA: The Biomechanics and Motor Control of Human Gait: Normal, Elderly and Pathological. Waterloo: University of Waterloo Press, 1991. 160. Winter DA, Patla AE, Frank JS, Walt SE: Biomechanical walking patterns in the fit and healthy elderly. Phys Ther 1990; 70: 340–347. 161. Winter DA, Quanbury AO, Reimer GD: Analysis of instantaneous energy of normal gait. J Biomech 1976; 9: 253–257. 162. Winter DA, Robertson DGE: Joint torque and energy patterns in normal gait. Biol Cybern 1978; 29: 137–142. 163. Winter DA, Yack HJ: EMG profiles during normal human walking: stride-to-stride and inter-subject variability. Electroencephalogr Clin Neurophysiol 1987; 67: 402–411. 164. Woodburn J, Turner DE, Helliwell PS, Barker S: A preliminary study determining the feasibility of electromagnetic tracking for kinematics at the ankle joint complex. Rheumatology (Oxford) 1999; 38: 1260–1268. 165. Woolley SM, Sigg J, Commager J: Comparison of change in level walking activities in three groups of elderly individuals. Gait Posture 1995; 3: 81. 166. Wu G, Ladin Z: Limitations of quasi-static estimation of human joint loading during locomotion. Med Biol Eng Comput 1996; 34: 472–476.