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Gait Analysis and Training of People With Limb Loss

Chapter 19 

GAIT ANALYSIS AND TRAINING OF 

PEOPLE WITH LIMB LOSS

JASON M. WILKEN, PhD, PT*; and RAUL MARIN, MD†

INTRODUCTION

TERMINOLOGY OF GAIT ASSESSMENT

OBSERVATIONAL GAIT ASSESSMENT

COMPUTERIZED GAIT ASSESSMENT

ABLE-BODIED (“NORMAL”) GAIT

GAIT DEVIATIONS IN THE LOWER LIMB AMPUTEE

STANDARDIZED GAIT ASSESSMENT

TRANSTIBIAL AMPUTEE GAIT

TRANSFEMORAL AMPUTEE GAIT

SPECIAL CONSIDERATIONS

GAIT LABORATORY DEPENDENT THERAPEUTIC INTERVENTIONS

SUMMARY

*Director, Military Performance Laboratory, Department of Orthopaedics and Rehabilitation, Center for the Intrepid, Brooke Army Medical Center, 3851 

Roger Brooke Drive, Fort Sam Houston, Texas 78234

†Colonel (Retired), Medical Corps, US Army; Physician, Center for the Intrepid, Brooke Army Medical Center, 3851 Roger Brooke Drive, Fort Sam 

Houston, Texas 78234; formerly, Medical Director, GaitLab, Physical Medicine and Rehabilitation Teaching Staff, and Chair, Internal Review Board, 

Department of Orthopaedics and Rehabilitation, Walter Reed Army Medical Center, 6900 Georgia Avenue, NW, Washington, DC
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Care of the Combat Amputee

INTRODUCTION

tion, in turn, provides an opportunity to develop corrective strategies designed to enhance the efficiency, 

comfort, and cosmesis of amputee gait. Attainment of 

“normality” is not an end in itself, but rather a guide 

toward the optimization of function.

This chapter reviews the terminology of gait assessment and discusses both observational and 

computerized gait assessment. The authors’ current 

understanding of able-bodied gait and documentation of gait performance observed in current military 

amputees is presented. Also included is a discussion 

of several special considerations about amputee gait 

and how gait laboratory assessment translates into 

effective therapeutic interventions. 

The systematic study of human ambulation began 

in the mid-1800s with the appearance of photography 

and observation-based studies.1 Subsequently, kinetic 

(forces)2 and kinematic (joint movement)3 studies began 

near the onset of World War II. The large number of amputees resulting from this conflict became a catalyst for 

the use of technology for the assessment of amputee gait. 

Inman and his colleagues introduced the use of movies 

in the coronal, sagittal, and transverse planes, as well as 

the use of force plates and electromyography (EMG) into 

the evaluation of able-bodied and amputee gait.4 

The intent of amputee gait assessment is the identification of gait parameters that deviate or differ from 

able-bodied gait (ie, from “normal”). This identificaTERMINOLOGY OF GAIT ASSESSMENT

A clear understanding of the terminology used in 

the description and evaluation of gait is essential prior 

to the assessment of patient performance. A clear or 

well-defined “common language” among the various 

disciplines forms the foundation on which a systematic 

and reliable assessment is built. Such assessment, in 

turn, leads to the effective planning and execution of 

interventions. Furthermore, common language facilitates a systematic tracking of progress because it allows 

for consistent pre- and postintervention comparison 

and monitoring.

The gait cycle consists of one stride that is divided 

into two periods known as stance and swing periods. 

During stance, the primary responsibility of the limb 

is to support the superincumbent body weight, while 

during swing, the task of limb advancement is accomplished. Figure 19-1 provides a visual representation 

of the gait cycle.5 Note that double support forms 20% 

to 30% of the gait cycle and that the terminology for 

describing the phases of gait has changed in recent 

years. To account for the frequent absence of heel 

strike in pathologic gait and to allow for a more consistent framework for discussing gait deviations, the 

classic terminology of heel strike, foot flat, midstance, 

heel-off, toe-off; and early, mid, and terminal swing 

is used less frequently. Rather, initial contact, loading 

response, midstance, terminal stance, and pre-swing 

are commonly used to describe the stance phase. Initial, middle, and late swing are used to characterize 

the swing phase.

One of the easiest temporal and spatial parameters 

to measure quantitatively is walking speed (meters/

second). This parameter is critical because velocity 

strongly influences the loading, motion, and alignment of the different joints throughout the gait cycle. 

Furthermore, each individual selects his or her natural 

rate of walking so that the speed selected minimizes 

energy expenditure.4 When the gait efficiency of an 

individual is plotted against various walking speeds 

ranging from very slow to very fast, a parabolic curve 

is displayed demonstrating that speeds slower or faster 

than the self-selected walking velocity (SSWV) will 

increase energy expenditure.6

There are several other temporal and spatial measurements that are used in the assessment of gait. Step 

length is the distance between a person’s heels during 

the double support phase of gait. Stride length is the 

distance traveled by the heel from initial contact of one 

foot to the subsequent initial contact of the same foot. 

One stride length contains two step lengths. The number of steps that occur within a given period of time 

(typically per minute) is known as cadence. Cadence 

can be multiplied by the individual’s stride length to 

determine walking velocity.5

Bipedal human locomotion is the result of evolutionary changes that allowed—among other 

things—the freeing of the upper limbs. This evolutionary change from quadruped to biped locomotion 

involved changes in the action of forces across levers 

responsible for locomotion. Torque from application 

of force distant to the axis of rotation that results in 

motion about the axis of rotation (if not opposed) is 

referred to in this chapter as a joint moment or simply 

moment. The convention of internal joint moments 

will be used throughout this chapter. Internal joint 

moments are commonly used because they are associated with the internal (typically muscular) forces that 

are necessary to produce or resist rotational motion 

at a given joint. These forces are in direct response to 

external moments, which are generated by the forces 
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Gait Analysis and Training of People With Limb Loss

Figure 19-1. Time dimensions of the gait cycle.

Reproduced with permission from: Rose J, Gamble JG, Eds. Human Walking. Baltimore, Md: Lippincott Williams & Wilkins; 

2005:26.

Right

heel

contact

Left

heel

contact

Left

heel

contact

Right

heel

contact

Left

toeoff

Right

heel

contact

Left

toeoff

Right

toeoff

0% 50% 100%

Time, percent of cycle

Double

support

Double

support

Double

support R. Stance support L. Stance support

R. Stance phase R. Swing phase

L. Swing phase L. Stance phase

Cycle (stride) duration

Right

heel

contact

Cycle length

(stride)

R. Step

length

R. Step

length

L. Step

length
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Care of the Combat Amputee

that push against the body (eg, ground reaction force). 

In standard reporting of gait laboratory data the term 

moment refers to net or resultant joint internal moment or torque produced by muscular, ligamentous, 

or bony force.

To the clinician performing an observational gait 

assessment, the temporal and spatial components 

of gait are—at most—approximations of observed 

deviations from what an expected symmetric gait 

should be. Computerized gait assessment, however, 

provides for precise quantifiable measurements of 

these variables.

OBSERVATIONAL GAIT ASSESSMENT

In the military setting a wide range of tools is 

available to assess amputee gait performance. These 

tools range from the brief informal visual assessment 

commonly used in the clinic to the application of 

advanced motion capture technologies. Selection of 

the most appropriate assessment approach is based 

primarily on the desired outcome of that particular 

evaluation session, which is determined using the 

patient’s and therapist’s rehabilitation goals, as well 

as the standardized assessment protocols at each 

facility. Observational assessment is—by its very 

nature—prone to inaccuracies. When a patient is 

walking, the speed at which motion occurs is commonly faster than what can be processed by the 

observing clinician. Intuitively, one would think 

that this limitation can be addressed by the presence 

of multiple observers under the belief that what 

one observer may miss may be caught by another 

observer. However, Krebs et al noted that single observer observational gait analysis is more consistent 

than multiple observer gait analysis and that this 

consistency could be increased for both methods by 

video recording, particularly if the replayed video 

was in slow motion.7

If computerized gait analysis is used as the gold 

standard by which observational gait analysis is 

to be compared, then it is apparent that—although 

convenient—observational gait analysis lacks adequate sensitivity. Saleh and Murdoch demonstrated 

that visual observers recorded only 22% of the gait 

laboratory-determined gait deviations with the measurement system detecting 3.4 times more deviations 

than visual observation alone.8

Although the above discussion makes a strong 

case for the application of computerized gait analysis, 

the complexity, cost, and limited availability of this 

methodology significantly hampers its clinical utility. Thus, the average clinician is left with his or her 

observations to assess the degree of pathology and 

the intervention required to address that pathology. 

A report by Gage in 2004 recommends that a systematic and well-organized observational gait analysis 

recording form be used for all such evaluations.5

The Rancho Los Amigos full body observational gait 

analysis form is perhaps the most well-known tool to 

accomplish this evaluation.9

COMPUTERIZED GAIT ASSESSMENT

Gait laboratory assessment in the military setting 

typically consists of video-based observational analysis in combination with quantitative assessment of 

temporal-spatial, kinematic, and kinetic parameters. 

This assessment differs significantly from the typical 

evaluation performed in the clinic in that it provides 

quantitative information that can be used to identify 

gait deviations, provide insight into joint loading and 

torques, and objectively track progress through preand postintervention comparisons. Computerized gait 

analysis is the standard of care for military amputees 

and is part of a regular assessment plan. Clinical experience suggests that the best results are obtained in 

cases where the gait pattern is complex or deviations 

are subtle, the altered gait pattern is believed to be the 

source of decreased function or pain, and the patient 

is highly motivated and follows the resulting treatment plan.

Data Collection

To allow data sharing, facilitate troubleshooting, 

and support collaborative research projects, the Military Advanced Training Center and the Center for the 

Intrepid military amputee patient care sites use similar 

assessment tools for quantifying gait performance.

To provide an accurate assessment of amputee gait 

kinematics (motion), both sites use passive markerbased optoelectronic motion tracking systems. Passive 

tracking systems rely on the placement of small (9 mm) 

reflective markers on the patient’s body to accurately 

track joint and segment positions in three-dimensional 

space. To independently monitor the three-dimensional 

motion of each segment, a minimum of three tracking 

markers is placed on each segment. These tracking 

markers are supplemented with additional markers 

that are used to identify bony landmarks and define 
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Gait Analysis and Training of People With Limb Loss

the orientation of each segment (eg, thigh). The total 

number of markers used during a session can range 

from 6 to 70, depending on the type of assessment 

being performed.

Between 12 and 26 motion capture cameras emit 

infrared light that is reflected off the surface of the 

tracking markers and then recorded on the camera’s 

image sensor. Image processing is performed within 

the camera to identify the centroid—or center—of the 

marker as a two-dimensional coordinate (vertical and 

medial-lateral position). The information from each 

camera is then sent to a central data collection computer. By combining two-dimensional images from the 

cameras, the system can triangulate the three-dimensional position of the marker with an accuracy of 0.7 

millimeters. These marker position data are then used 

to quantify the magnitude and direction of motion that 

occur at individual joints and are presented in graphical format for interpretation by the clinical staff.

Force plates are used in all amputee center gait 

laboratories to assess gait kinetics (or forces). Typically 

eight force plates are embedded in the main walkway 

allowing the ability to measure location, direction, and 

magnitude of force applied at the foot floor interface. 

Although ground reaction force data are useful as 

an independent assessment tool, the force data are 

typically combined with marker data to calculate joint 

moments as well as power generation and absorption. 

Insight into the compensatory responses exhibited by 

amputees is provided using an approach called inverse 

dynamics. When using this method, joint position, 

segment motion, mass properties, orientation, and 

external forces acting on the segment are used to estimate lower extremity joint reaction forces, moments, 

and powers. Although not providing direct information regarding activation of individual muscles, the 

approach provides insight into the net effect of muscle 

activation at each joint of interest.

Because of the large effect of walking velocity on 

gait performance, gait kinetics in particular, patients 

are asked to walk along the laboratory walkway at both 

controlled and SSWV during standard clinical assessment. Collecting gait data at controlled speeds allows 

the opportunity to determine whether deviations from 

normal self-selected gait are a result of the amputation 

or simply the result of changes in walking velocity as 

compared to uninjured individuals.

Energy Expenditure

Metabolic cost assessment, another assessment technique that can provide insight into gait performance, is 

commonly used to assess the effectiveness of prosthetic 

components (as well as training interventions). Metabolic cost assessment is typically performed as patients 

ambulate on a treadmill while walking velocity is 

incrementally increased. Patients are typically tested 

at three to six controlled walking velocities approximately centered about the individual’s SSWV. Such 

assessment allows the ability to determine the patient’s 

gait efficiency and his or her ability to accommodate 

ambulation at multiple speeds. 

Electromyography

EMG can be used to provide insight into the timing, magnitude, and patterns of muscular activation 

used by amputees. Use of surface electrodes allows 

the ability to quantify muscle activation while minimizing the risk of discomfort and infection that can 

be associated with indwelling electrodes. Although 

challenges such as appropriate normalization of EMG 

data and reduced access to residual limb musculature 

must be overcome, the assessment of muscle activation provides insight that other approaches cannot. 

Unlike the inverse dynamic approach that is unable 

to account for co-contraction and force production 

by passive structures (bony or ligamentous restraint), 

EMG allows the ability to directly record the timing 

and magnitude of muscle activation.

ABLE-BODIED (“NORMAL”) GAIT

In 1985 Perry described four essential requirements for normal gait that—when absent in part or in 

whole—produce deviations from the norm.10 These 

requirements are (1) symmetric or proportionate step 

length, (2) stance stability, (3) swing clearance, and 

(4) adequate foot position before initial contact. Gage 

added energy conservation as a fifth essential requirement in 1991.11

Energy conservation during ambulation is multifactorial in nature. The body is designed to be able to 

take advantage of the ground reaction force for joint 

stability and efficiency. Efficiency is also increased by 

the predominant use of eccentric muscle contractions 

during movement. Furthermore, when the less efficient 

concentric contractions occur, the muscle begins from 

a stretched or elongated position to allow the recovery 

of kinetic energy stored in elastic connective tissues. 

Energy conservation also occurs through muscles 

crossing two joints that serve the function of linking 

and transferring forces across body segments. 
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Care of the Combat Amputee

The phases of the gait cycle depicted in Figure 19-1 

are helpful in understanding the sequence of events 

that is characteristic of a normal gait cycle. Human 

walking can be described as a process of alternating 

acceleration and deceleration of the center of mass, 

with the bulk of energy utilization occurring at the 

beginning and at the end of the stance periods. As such, 

during loading response the quadriceps and tibialis anterior muscles contract eccentrically to allow controlled 

knee flexion and ankle plantarflexion resulting in a 

net absorption of force and subsequent deceleration 

of the body. During midstance the momentum generated by the above described muscle actions takes the 

body’s center of gravity over the weight-bearing leg 

as the contralateral leg swings like a pendulum with 

minimal energy utilization. During terminal stance 

and pre-swing, there is ipsilateral force generation 

and acceleration via the concentric contraction of the 

gastroc-soleus complex and similar response of the 

gluteus maximus in the contralateral side.

Evolutionary changes in human anatomy and 

biomechanics have ensured that human ambulation 

occurs with the greatest efficiency at self-selected 

walking speeds. In 1953 Saunders et al proposed the 

six determinants of gait as a way of explaining how 

the body’s displacement of the center of mass within 

a 5-cm sinusoidal pattern of horizontal and vertical 

displacement conserves energy expenditures during 

walking.12 Saunders et al’s six determinants of gait 

follow: 

1. Pelvic tilt: in the frontal plane, during stance 

the weight-bearing gluteus medius allows a 

drop of the contralateral swing side iliac crest, 

thus decreasing the rise in the center of mass 

during single limb support.

2. Pelvic rotation: in the horizontal plane during 

double limb support, pelvic rotaion allows 

increased step length while minimizing the 

drop of the center of mass.

3. Knee flexion at loading response and during 

stance: reduces the rise of the center of mass 

as the body transitions over the stance limb.

4. Ankle dorsiflexion at initial contact: effectively elongates the limb allowing a smooth 

transition onto the stance limb preventing a 

drop of the center of mass.

5. Ankle plantarflexion at terminal stance/preswing: effectively elongates the limb allowing 

a smooth transition off of the stance limb by 

preventing a drop of the center of mass.

6. Side-to-side pelvic displacement from stance 

to stance: shifts the center of mass from one 

weight-bearing leg to the other.

For decades, these six determinants of gait were 

believed to effectively characterize the major factors 

that serve to minimize energy expenditure in human 

ambulation under the premise that they minimized 

the excursion or displacement of the center of gravity. 

Although evidence supports the association between 

vertical displacement of the center of mass and energy 

consumption, other studies have questioned the premise that the six determinants of gait (in combination 

or independently) are the sole factors responsible for 

controlling the center of mass displacement and thus 

energy expenditure.13

As early as 1983, in a study measuring mechanical 

cost at three different walking speeds (slow, self-selected, fast), Winter found that energy cost decreases as 

knee extension increases during stance with a significant positive correlation between energy cost and the 

degree of knee flexion.14 In a series of articles, Kerrigan 

and her co-investigators found that heel rise from foot 

flat to pre-swing significantly raises the center of mass 

when it is at its lowest point during the gait cycle with 

the end result of reducing its overall displacement. 

They also found that pelvic rotation contributes no 

more than 12% of the total reduction of the vertical 

displacement of the center of mass.15,16 Likewise, other 

investigators have documented that although knee 

flexion reduces the center of mass vertical excursion, 

it significantly increases the energy costs, and that heel 

rise from foot-flat to pre-swing is responsible for twothirds of the reduction in the center of mass vertical 

excursion.17-19

It appears then that the six determinants do not 

have the same influence in reducing the center of 

mass excursion and that perhaps other as-of-yet to be 

identified determinants must be contributing to the 

energy cost savings resulting from the minimization 

of the center of mass displacement.

GAIT DEVIATIONS IN THE LOWER LIMB AMPUTEE

As in the civilian setting, lower limb military amputees typically demonstrate characteristic gait deviations as a result of the partial loss of lower limb function. Gait deviations observed in the amputee patient 

are commonly attributed to the following:

	 •	 loss of active torque generation;

	 •	 loss of somatosensory feedback and limb position awareness; 

	 •	 additional degrees of freedom added by the

mobile interface between the residual limb 
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Gait Analysis and Training of People With Limb Loss

and socket; 

	 •	 pain;

	 •	 limitations of current prosthetic devices (foot,

socket, and knee if applicable); and 

	 •	 functional impairments in the contralateral

limb. 

Fortunately, with adequate training, prosthetic care, 

and patient motivation, these factors can be minimized 

or compensated for, yielding a very functional gait 

pattern that is nearly indistinguishable from that of 

uninjured individuals.

Many factors influence the maximal gait performance 

patients attain before discharge from treatment. In military amputees, the frequent presence of significant injury 

to multiple areas of the body plays an important role in 

determining the maximal gait performance attained after 

completion of therapy. Patients often experience significant bony and soft tissue injury to the nonamputated limb 

at the time of injury, which can result in a nonamputated 

side that is less functional than the amputated side. In 

some instances the amputated side is referred to as the 

“good” or more functional side. Because other injuries 

vary widely in their impact on the patient’s overall gait 

performance, the current discussion will focus on what 

is currently deemed the maximal gait performance exhibited by the highest functioning patients and then on 

commonly observed deviations from that gait pattern.

When assessing gait performance it is important to 

consider the potential long-term impact of individual 

gait deviations. Although the patients and clinicians 

may wish to minimize all gait deviations until they are 

not readily noticeable, efforts should first be focused 

on facilitating the most functional and efficient gait 

while identifying and addressing the source of gait 

deviations that are most likely to produce long-term 

morbidity such as osteoarthritis, low back pain, and 

other commonly observed problems. 

STANDARDIZED GAIT ASSESSMENT

To provide a consistent and thorough assessment of 

amputee gait, a standardized approach is used for the 

interpretation of gait performance. A comprehensive 

assessment typically begins with the assessment of 

temporal-spatial parameters such as walking velocity, 

step length, step width, stride length, and other gait 

parameters that allow the rapid detection of gross 

asymmetries and other deviations from normal. 

Information such as walking velocity is particularly 

valuable when interpreting kinematic and kinetic data 

that are influenced by walking velocity. Kinematic data 

are then assessed using what is commonly known as a 

“bottom-up” approach, beginning with the examination of ankle and foot motion and progressing up the 

kinematic chain. These measures are compared with 

that of uninjured individuals and high functioning 

patients that are thought to demonstrate maximal or 

“optimal” gait performance. After completing the assessment of kinematic data, torque production at the 

ankle, knee, and hip is assessed in a similar manner, 

beginning at the ankle and ending at the hips. This 

is followed by assessment of joint power data before 

providing a comprehensive assessment of the patient’s 

overall gait pattern and final report generation.

TRANSTIBIAL AMPUTEE GAIT 

Due to their relatively young age and high fitness 

level, service members who have experienced transtibial amputation with adequate residual limb length 

and minimal contralateral involvement are typically 

able to achieve a gait pattern that is indistinguishable 

from uninjured individuals. The highest functioning 

individuals demonstrate gait patterns that are fluid, 

functional, and approximate the gait of uninjured 

individuals in terms of gait kinematics and kinetics: 

their gait patterns are used as the benchmark against 

which other amputees are compared. For this reason, 

deviations from normal gait that are demonstrated by 

patients with maximal or “optimal” gait performance 

will be presented before discussing additional deviations that are addressed with training or prosthetic 

management. Despite the extensive study of amputee 

gait, the “optimal” gait pattern has yet to be defined. 

Initial Contact through Midstance

For most transtibial amputees, the stance phase 

begins with the lower limb positioned in a manner 

nearly identical to that of uninjured individuals. The 

ankle is in a neutral position, the knee is nearly fully 

extended, and the hip is flexed to between 25 degrees 

and 35 degrees. Differences from normal gait are, 

however, observed as soon as the patient begins to 

bear weight on and control motion of the involved 

limb and prosthesis.

Ankle Kinematics and Kinetics

State-of-the-art prosthetic feet currently provided 

to military amputees function—mechanically speaking—in a manner that is significantly different from 
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Care of the Combat Amputee

the physiologic ankle. Shock absorption and “ankle” 

motion after initial contact are achieved by compression of the heel component rather than through true 

plantarflexion rotation about a physiologic ankle joint. 

Observed motion is, therefore, primarily an artifact 

of proximal displacement of the heel relative to the 

forefoot resulting in rotation of the shoe relative to 

the shank. However, meaningful information can be 

gained from the interpretation of “ankle” kinematic 

and kinetic data as will be established during presentation of specific gait deviations associated with 

prosthetic alignment, foot category, and training.

During loading response most high functioning 

patients are able to reproduce sagittal plane ankle 

kinematics that are indistinguishable from uninjured 

individuals. Although they produce normal kinematics, most patients demonstrate dorsiflexion torque that 

is increased in both duration and magnitude.20,21 This 

is due—in part—to the frequently observed increase in 

double limb support time, which is believed to impart 

additional stability, and compression of the prosthetic 

heel, which results in patients spending a significant 

amount of time on the heel. Once patients reach a stable 

foot flat position and transition off the heel component, 

ankle kinematics and kinetics are nearly identical to 

that of uninjured individuals until pre-swing. As loading shifts from the heel to the forefoot due to forward 

rotation of the tibia associated with progression of 

the trunk over the stance limb, gradual dorsiflexion 

motion and an increasing plantarflexor moment are 

observed. Resulting sagittal plane ankle kinematics 

and kinetics (moment and power) during midstance 

are commonly normal with respect to both pattern and 

peak magnitude. 

Knee Kinematics and Kinetics

Consistent with the broader amputee population, high functioning patients typically display 

decreased knee flexion from loading response until 

pre-swing.20,22,23 Although knee flexion during the 

stance phase is present, it is both decreased and 

delayed relative to nonamputee gait.22 Although 

nonamputee gait is commonly seen as ideal, a 

convincing case has not been made to indicate that 

decreased knee flexion during stance phase is detrimental and not an effective compensatory strategy. 

There are several potential explanations why patients 

choose to decrease knee flexion. First, this altered 

gait pattern may compensate for the functional loss 

of leg length because of the distal translation of the 

residual limb within the socket and compression of 

the prosthetic foot and shock absorber (if present).24-29

Second, decreased knee flexion during midstance 

may assist in the gradual storage of energy in the 

prosthetic forefoot that can be released for propulsion. With an extended knee the body moves as an 

inverted pendulum. Forward movement of the body 

over the fixed foot and rotation of the extended leg 

result in gradual dorsiflexion of the prosthetic foot 

and storage of energy that can be returned during 

pre-swing. In this manner, amputees are also able 

to effectively utilize the rollover characteristics of 

the prosthetic foot.30,31 Third, co-contraction may be 

used as a strategy to stabilize the mobile bone–socket 

interface and decrease the degrees of freedom of the 

residual limb.24,26-29 Increased activation of the quadriceps muscle, despite an extended knee position, has 

been previously identified as a strategy to stabilize 

the knee.21,22 Although this strategy likely serves to 

control the knee, it may also be in response to knee 

flexor torque produced by activation of the biarticular 

gastrocnemius muscle and an overall stiffening of 

the limb. Activation has been identified as a strategy 

to stabilize the bone–socket interface by increasing 

the stiffness of the soft tissue in the residual limb.32

Although sagittal plane kinematics of the knee is 

altered, as previously reported net joint torques are 

minimally affected.22

Hip Kinematics and Kinetics 

Sagittal plane kinematics and kinetics of the hip 

have been identified as the primary source of compensation in civilian transtibial amputees, but are 

relatively unaffected in high functioning military 

amputees.20,21 Hip flexion is occasionally increased at 

initial contact and toe-off but this does not significantly 

alter joint torques or powers. The increased hamstring 

activation associated with reported co-contraction at 

the knee and other reported changes in muscle activation are either not present or have little effect on the 

sagittal plane kinematics and net sagittal plane hip moments for military amputees. In the highest functioning 

transtibial amputees the most frequent deviation in 

hip kinematics is an occasional increase in hip flexion 

at initial contact.21

Trunk and Pelvic Kinematics

A unique characteristic of very high functioning 

military transtibial amputees is the ability to produce 

frontal plane trunk and pelvic kinematics that are 

similar to that of uninjured individuals. Deviations 

in frontal plane kinematics are documented in other 

populations but are typically most evident early in the 

rehabilitation process for military amputees.33 This deviation is apparently reduced through the combination 
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Gait Analysis and Training of People With Limb Loss

of extensive training using manual and verbal feedback 

in combination with proper prosthetic management. 

Although normal kinematics can be replicated in high 

functioning transtibial amputees, as a general rule 

patients demonstrate increased frontal plane motion 

at the trunk and pelvis during gait. 

Pre-Swing

Gait kinematics and kinetics during pre-swing 

and swing in the military population are similar to 

those of civilian amputees with a few key exceptions. For all amputees using dynamic energy storing and return feet, the peak dorsiflexion angle that 

is achieved is determined by a combination of foot 

stiffness, alignment, and the ankle joint torque produced by the patient due to loading of the trail limb. 

As observed in previous studies, with appropriate 

prosthetic management patients are able to attain 

normal plantarflexor torque and motion at terminal 

stance. Patients demonstrate a remarkable ability to 

control the rate of energy returned from the foot and 

as a result are able to achieve nearly normal power 

output at the ankle. This observation is in sharp 

contrast to literature indicating “marked reduction 

in energy generating capabilities compared to the 

normal gastrocnemius-soleus.”20,21,34-36 Although the 

exact mechanism by which normal ankle powers are 

achieved is unknown, current explanations focus on 

training to sustain loading of the trail leg until just 

before the foot is lifted from the floor. The ability to 

produce normal ankle power is particularly impressive given that current prosthetic feet only return to 

the neutral alignment position (typically in slight 

dorsiflexion) when unloaded.

Swing

With the exception of ankle kinematics, few differences exist between the swing phase kinematics of 

high functioning transtibial amputees and uninjured 

individuals. Toe clearance (1–2 cm) is the most important aspect of the swing phase in transtibial amputees. 

The loss of active control of ankle motion prevents patients from using dorsiflexion during midswing to aid 

in clearing the toe. This is particularly relevant when 

considering the effective increase in leg length that occurs due to distal translation of the socket relative to 

the residual limb during swing.37 Although microprocessor controlled prosthetic foot and ankle systems are 

available to aid with toe clearance, they have not been 

widely accepted by the patients. Instead patients use 

nearly imperceptible compensatory changes at more 

proximal joints that allow adequate clearance.

Additional Kinematic and Kinetic Deviations in 

the Transtibial Amputee 

Deviations in ankle kinematics and kinetics are 

most commonly observed early in the rehabilitation 

process as patients are beginning to walk or are accommodating to new prosthetic components. In the 

authors’ experience, ankle kinematics and kinetics 

are primarily influenced by the stiffness (grade) of the 

prosthetic foot and training to control loading of the 

prosthetic foot.

Initial Contact Through Midstance

One of the most commonly observed deviations in 

transtibial military amputees includes ankle plantarflexion motion and ankle dorsiflexion torques that are 

increased in duration and magnitude relative to other 

transtibial amputees.20,21 These deviations are primarily 

attributed to a prosthetic foot that is too soft or too excessive and poorly controlled loading early in stance phase. 

In the case of a soft foot, the heel undergoes significant 

compression resulting in increased time in loading response. Rather than gradually moving from the heel during loading response to foot flat, this transition is delayed 

as the patient compresses the prosthetic heel. The aforementioned deviations in ankle kinematics and kinetics are 

typically accompanied by deviations at more proximal 

joints. Knee flexion is decreased or nearly absent resulting in a gait pattern similar to that of high functioning 

transfemoral amputees. As the heel compresses, patients 

lean toward the stance limb resulting in frontal plane 

rotation of the trunk and pelvis toward the stance limb.

In contrast, some individuals demonstrate an opposite response in which ankle plantarflexion motion and 

ankle dorsiflexion torques are decreased in duration 

and magnitude.21 This is typically observed in patients 

with a prosthetic foot that is too stiff and/or in cases 

where patients are experiencing anterior-distal pain or 

have a poor tolerance for loading early in stance phase. 

In both instances, patients tend to move quickly to a 

foot flat position following initial contact. In the case 

of a foot that is too stiff, it appears as though patients 

avoid the unstable position of balancing on a firm heel 

and instead quickly move to a more stable foot flat 

position. For patients with a painful residual limb, a 

compensatory response of a foot flat landing and decreased time on the heel is reasonable given the current 

understanding of limb–socket dynamics. At initial contact a significant anterior and inferiorly directed force is 

applied at the bone–socket interface resulting in distal 

translation of the limb within the socket and loading 

along the anterior aspect of the socket. By landing in a 

foot flat position, patients are able to decrease the shear 
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Care of the Combat Amputee

force during loading response and direct force along 

the axis of the residual limb into the soft tissue bulk 

instead of focusing it on the anterior-distal end. Due 

to the inability to actively control ankle orientation to 

bring the forefoot to the floor, patients use increased 

knee and hip flexion motion to lower the foot to the 

floor in a controlled manner.21 Knee extensor torque is 

not, however, typically increased in military amputees 

because of efforts to minimize anterior shear forces in 

the painful patient.

Pre-Swing

Improper stiffness (grade), alignment of the prosthetic foot, or poor muscular control also results in 

gait deviations during pre-swing. As previously 

mentioned, the alignment and stiffness of the foot can 

significantly influence the peak dorsiflexion angle and 

response of the foot during pre-swing. A forefoot that 

is too soft or is aligned in excessive dorsiflexion can 

allow excessive peak dorsiflexion motion and is commonly accompanied by patient reports of insufficient 

foot length or energy return. Rather than exhibiting 

increased peak dorsiflexion, patients commonly accommodate by decreasing loading of the trail limb, 

which has the negative consequence of decreased 

plantarflexor power. Although compensations such 

as increased knee and hip extension are additional 

strategies that may be used to accommodate a prosthetic foot that is too soft, such responses are not 

consistently observed.

A forefoot that is too stiff or is aligned in excessive 

plantarflexion does not allow sufficient dorsiflexion 

motion and is accompanied by increased knee extension during mid and late stance.21 This is due to 

increased resistance to forward rotation of the tibia 

over the foot. Rather than increasing the loading on 

the forefoot to attempt to deflect the foot, patients 

typically compensate by extending the knee. As a 

result, dorsiflexion motion is decreased, which leads 

to decreased peak plantarflexor power as the foot is 

unloaded at terminal stance.21

TRANSFEMORAL AMPUTEE GAIT

The combined loss of ankle and knee function and 

a dynamic bone–socket interface results in more pronounced gait deviations in patients that have undergone transfemoral amputation or knee disarticulation 

procedures. Although military transfemoral amputees 

have access to the latest prosthetic technologies and 

training, they typically have gait deviations that are 

readily apparent to an untrained observer. Appropriate 

prosthetic and physical therapy management is essential to allow restoration of a functional, efficient gait 

pattern. Contracture prevention, strength training, and 

appropriate socket fit are critical to provide patients the 

range of motion, strength, and control of the prosthesis 

necessary to produce a good gait pattern.

Initial Contact Through Midstance

Ankle Kinematic and Kinetic Data

Sagittal plane ankle kinematics and kinetics of the 

high functioning transfemoral amputee closely mimic 

those of the transtibial amputee. They typically demonstrate near normal sagittal plane ankle kinematics 

but exhibit the frequently observed pronounced dorsiflexion moment at loading response.38-40

Knee Flexion Kinematic and Kinetic Data

Sagittal plane knee kinematics differs significantly 

from that of uninjured individuals. Like most civilian 

amputees, military transfemoral amputees have little 

if any knee flexion and exhibit knee flexor torques 

throughout the stance phase.

38-41 This is primarily 

attributed to a training philosophy that encourages 

patients to utilize hip extensor activation to stabilize 

the prosthetic knee by pushing it into full extension 

during stance. There are three perceived benefits of 

this training. First, by not relying on the 10 degrees 

to 15 degrees of stance flexion allowed by most prosthetic knees, there is a greater margin for error when 

performing challenging tasks such as ambulating on 

uneven terrain. If patients land on an object that serves 

to push their knee into flexion, they are already actively 

creating stability rather than relying on the response of 

the prosthesis. If the knee begins to buckle, the muscles 

are already tensioned leading to a faster response 

time and decreased likelihood of falling. Second, it 

encourages activation of hip extensor musculature for 

propulsion decreasing demand on the contralateral 

limb and easing the transition to high-level activities, 

such as running, during which hip extensor torque is 

a primary means of propulsion. Third, the repeated 

activation of residual limb musculature that is encouraged with this approach is thought to reduce muscle 

atrophy (maintain muscle volume) and provide a more 

rigid limb–socket interface.

Although stance flexion was thought to play an 

important role in minimizing center of mass excursion 

and therefore metabolic cost, its value has recently been 

questioned. More recently the focus on stance flexion 
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Gait Analysis and Training of People With Limb Loss

has revolved around its role as a shock absorbing 

mechanism. Although likely useful for some patients, 

in the authors’ experience, stance flexion is not necessary because of adequate shock absorption provided 

by the inclusion of rotation and shock absorbers, 

deflection of the prosthetic heel, and the bone–socket 

displacement that occurs in current prosthetic socket 

designs.30 This ability to adequately absorb impact 

forces is clearly evident in ground reaction force curves 

that typically lack the pronounced vertical impact 

peaks observed in normal gait.9

Hip Kinematic and Kinetic Data

As a result of the closed kinematic chain present 

during stance, altered kinematics of the knee results 

in deviations in hip kinematics. Unlike normal gait 

in which the hip flexion angle changes little during 

the first 5% to 10% of the gait cycle, transfemoral 

amputees commonly demonstrate a more constant 

hip extension velocity (straight line on kinematic 

curve) from initial contact until peak hip extension. 

In uninjured individuals the minimal changes in hip 

flexion angles during early stance are the result of 

stance flexion moving the entire limb into a more 

flexed position resulting in a delayed transition toward extension at the hip. In transfemoral amputees, 

the absence of stance flexion results in a consistent 

hip extension velocity as the pelvis progresses over 

the fully extended leg. 

Trunk and Pelvic Kinematics

Typically only high functioning amputees with 

well-fitting sockets are able to minimize deviations in 

frontal, sagittal, and transverse plane pelvic and trunk 

kinematics that are commonly seen in civilians.33,42-44

It is unclear whether deviations in pelvic motion 

are the result of limitations in prosthetic technology, training, or both. Although the use of ischial 

containment socket designs allows freedom of hip 

and pelvic motion not otherwise attainable with an 

ischial weight-bearing quadrilateral socket, most 

patients demonstrate a compensated gluteus medius 

gait pattern. Rather than the contralateral side of the 

pelvis decreasing in height during loading response, 

the contralateral shoulder and pelvis elevate resulting 

in trunk and pelvis rotation toward the stance limb.42,44

This altered trunk and pelvic motion, which is commonly the most visually apparent gait deviation, is 

readily recognized by the untrained observer. Two 

primary causes of this increase in motion are believed 

to be (1) hip abductor weakness and the (2) mobility 

of the femur within the socket.30,45 By bringing the 

center of mass over the hip joint center the hip abductor moment is decreased, thus reducing the torque 

requirements of the gluteus medius. Similarly, frontal 

plane trunk and pelvic rotation toward the stance limb 

allows the patients to axially load through the socket 

rather than relying on compression of the lateral soft 

tissues or co-contraction of thigh muscles to stabilize 

the femur in the socket. 

Although some patients are able to achieve normal 

pelvic orientation (15 degrees anterior pelvic tilt) and 

excursions (approximately 5 degrees) throughout the 

gait cycle, a majority of patients demonstrate increased 

anterior pelvic tilt and sagittal plane motion.33 Commonly, patients move into a position of increased anterior pelvic tilt during late stance as they approximate 

hip end range of motion and then rapidly posteriorly 

tilt the pelvis to help initiate swing. Because of the 

perceived increased risk of low back pain associated 

with increased sagittal plane motion of the pelvis and 

lumbar spine, efforts are under way to more effectively 

train patients to decrease this motion.

Terminal Stance Heel Strike

Although transfemoral amputees approximate the 

nearly normal sagittal plane ankle motion and torque 

exhibited by transtibial amputees, subtle differences 

result in ankle joint power peaks that are consistently 

less than that exhibited by high functioning transtibial 

amputees.

38,40 It is, however, unclear whether this dif It is, however, unclear whether this difference is due to the loss of active control of the knee 

that alters loading of the foot, or the use of low profile 

feet that tend to be less dynamic.

During terminal stance, the knee joint angle rapidly 

moves from the extended position during midstance 

to approximate normal peak flexion angles during 

swing.

38,40 The rapid transition from extension into 

flexion is controlled by extensor torque produced by 

resistance within the prosthetic knee resulting in a 

burst of negative power that peaks around toe off.

Because it is the only source of active power 

generation in the involved limb, deviations are 

commonly seen in hip kinematics and kinetics. It is 

common for patients to demonstrate increased hip 

extension motion and increased hip flexor torque in 

late stance as compared to uninjured individuals. 

The most commonly observed deviation in sagittal 

plane pelvic motion is rapid motion into anterior 

pelvic tilt during terminal stance. Rather than isolating the motion to the hip joint they extend through 

the lumbar spine. Although functional, this motion 

pattern is discouraged because it is believed to 

contribute to the high prevalence of low back pain 

in amputees.
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Care of the Combat Amputee

SPECIAL CONSIDERATIONS

Energy Expenditure

The energy expenditure of an amputee is influenced 

by extrinsic factors such as prosthetic malalignment, 

leg length discrepancies, and the limb–socket interface 

mismatches previously discussed.

Intrinsic factors are those inherent to the amputee 

such as age and the presence of vascular disease. In addition, amputee specific gait deviations that result from 

the loss of sensory perception in the amputated limb 

(particularly proprioception), loss of the shock absorbing and propulsion properties of the foot and ankle (for 

transfemoral and transtibial amputees), and loss of the 

loading response at the knee (transfemoral amputee) 

lead to unavoidable deviations from the energy-efficient 

gait pattern of nonamputees. These losses can result in 

deviations that result in an increase in energy expenditure. As stated previously, although changes in the six 

determinants of gait have been used to explain these 

changes, some of these may actually have no influence 

and other factors yet to be identified may be more influential. Table 19-1 shows the percent increase in energy 

cost for the different levels of lower limb amputation.46

Note that these average values will increase when 

pathological deviations exist. The end result shows that 

to minimize these average energy cost increases, the 

amputee slows down the ambulation speed. Thus, just 

as in able-bodied individuals, the U-shaped curve seen 

when gait efficiency is plotted against various walking 

speeds is also seen in the amputee, although at a higher 

baseline. That is, the amputee consumes more energy 

when walking at slower or faster speeds than the SSWV 

with the energy expenditure of that SSWV being higher 

than in able-bodied individuals.47 As a general rule, appropriate prosthetic management of the amputee aims 

to keep the energy costs of ambulation at or below the 

values listed in Table 19-1 by minimizing previously 

discussed gait deviations.

In a review of the influence of prosthetic alignment 

and prosthetic components on energy expenditure, 

Schmalz et al found that malalignments affect transfemoral more than transtibial amputee gait.48 Furthermore, it was also found that there were no differences 

in energy expenditure when different prosthetic feet 

were worn by transtibial amputees and that microprocessor controlled knees reduced energy costs when 

compared to conventional hydraulic knees. Buckley 

et al also found that microprocessor knees tend to 

lower energy costs when compared to conventional 

pneumatic knees.49 Taylor et al found that although 

energy cost differences exist between these two prosthetic types at slower or SSWVs, there were energy cost 

savings when walking at higher speeds.50 Barth et al 

found no differences in metabolic costs among different feet in transtibial amputees although significant 

differences in temporal-spatial, kinetic, and kinematic 

parameters were found among the various feet tested.51

Furthermore, vascular amputees had higher metabolic 

costs than traumatic amputees. However, these studies had small sample sizes and lacked randomization 

or stringent controls. Large sample comprehensive 

randomized studies with stringent controls are needed 

to fully and conclusively assess the purported advantages of currently available advanced prosthetic 

components. 

Effect of Prosthetic Mass

As shown in Table 19-1, the average energy costs of 

prosthetic ambulation increase as the level of limb loss 

increases. This reality, with the evidence showing that 

adding weight to the lower extremities of nonamputees 

increases metabolic costs of walking, has triggered the 

development of lightweight prosthetic components 

over the past decade.52

The concerted effort to decrease prosthetic weight 

so as to decrease the metabolic costs of amputee 

gait, however, appears to have little scientific basis. 

Investigators have consistently demonstrated that 

no relationship exists between metabolic costs and 

prosthetic limb mass,53-56 while others have also shown 

that no relationships exist between prosthetic mass and 

temporal-spatial gait parameters such as stride length, 

stride frequency, or self-selected walking speed.55,57,58

Selles et al found that although variations in prosthetic mass do not seem to affect joint angles (ie, kinematics), changing limb mass does influence the forces 

being applied across the joints (ie, kinetics).59 Gitter et 

TABLE 19-1 

AVERAGE ENERGY COSTS OF PROSTHETIC 

AMBULATION

Percentage Increase

Unilateral transtibial 9%–28%

Unilateral transfemoral 40%–655%

Bilateral transtibial 41%–100%

Bilateral transfemoral 280%

Transtibial plus transfemoral 75%

Unilateral hip disarticulation 82%

Hemipelvectomy 125%
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Gait Analysis and Training of People With Limb Loss

al explained these changes in the kinetics of gait, however, in 1997 when they demonstrated that although 

an increase in the mechanical energy occurs across 

the prosthetic side hip joint during the pre-swing and 

swing phases of gait to accelerate the heavier prosthesis 

into swing, there is also a net absorption of energy during the terminal swing phase.57 Thus, the force needed 

to accelerate the heavier prosthesis during swing is 

counteracted by an equal degree of energy recovery 

as the limb decelerates during terminal swing with a 

net effect of zero loss of energy. 

To best serve the patient, the clinician should consider the location of the mass of the prosthesis rather 

than concentrate on the actual total weight of any 

given prosthesis. Lehmann et al demonstrated that 

the economy of gait decreases when the center of mass 

of the prosthesis is located distally.55 Furthermore, 

investigators have also determined that intermediate 

center of mass locations were the most energy-efficient 

locations for the prosthetic center of mass.52,56

Effect of Prosthetic Components

In an extensive review of the literature, van der 

Linde et al classified all the studies evaluating the effect of prosthetic components on amputee gait using 

established quality assessment criteria.60 The literature 

was divided into three categories:

1. A-level studies: studies meeting the most 

stringent criteria, thus defining these as studies of the highest quality.

2. B-level studies: studies partially meeting the 

selection criteria, thus defining these as studies of moderate quality. 

3. C-level studies: studies not meeting most of 

the criteria, thus defining them as studies of 

low-level quality.

A total of 40 of the 356 articles identified met the 

criteria warranting full assessment. These studies 

were then divided into prosthetic foot, prosthetic 

knee, prosthetic socket, and prosthetic mass focus 

studies. Within the prosthetic foot studies, one met 

the A-level criteria, 15 met the B-level criteria, and 5 

met the C-level criteria. As a whole, foot type did not 

influence the SSWV among individuals with traumatic 

transtibial and transfemoral amputation. There were 

few foot-specific effects on gait performance and the 

few foot-specific effects that were observed occurred

when energy-storing feet were compared with the 

Solid Ankle Cushion Heel foot. In general, although 

variability existed resulting from study group selection bias, the energy-storing feet appeared to provide 

for a higher SSWV. However, the evidence for a slight 

decrease in oxygen consumption is not convincing. 

Likewise, the evidence for a higher degree of patient 

satisfaction with the energy-storing feet is not fully 

convincing.

In this review, the authors also reported that as 

a whole, prosthetic knees offering advanced swing 

control (pneumatic, microprocessor-controlled, hydraulic) provide improved symmetry and gait velocity 

when compared to constant friction mechanical knees, 

particularly in active amputees. For the geriatric and 

more sedentary amputee, however, a locked knee 

with circumduction or hip hiking during swing may 

be more efficient because it provides improved knee 

stability required for this type of patient who is at risk 

of falls.60

Although van der Linde et al’s literature review used 

stringent criteria to select methodologically sound articles in their final analysis, only five manuscripts were 

included addressing the prosthetic knee, two of which 

were from the early 1980s, two from the mid 1990s, 

and one from 2000. Furthermore, their conclusions 

compared the pneumatic, hydraulic, and microprocessor knees versus the constant friction knee. Studies 

comparing microprocessor-controlled knees versus 

mechanical pneumatic or hydraulic knees and studies 

comparing various types of microprocessor-controlled 

knees were thus not included in their review.60

In an assessment of the C-Leg (Otto Bock Healthcare, Minneapolis, Minn) versus two hydraulic-based 

mechanical knees utilizing 10 amputees, Kastner et 

al reported that when using the C-Leg, amputees had 

the most normal kinematic parameters at various 

speeds and their walk was more efficient as demonstrated by having the fastest time in a 1,000-meter 

walk test.61 A more recent eight-subject study compared two microprocessor-controlled knees (C-Leg; 

RHEO KNEE, Ossur, Aliso Viejo, Calif) and a commonly used hydraulic mechanical knee (Mauch SNS, 

Ossur, Aliso Viejo, Calif).38 The outcome measures 

included (a) metabolic cost, (b) kinetic and kinematic 

data, and (c) EMG and accelerometer data. The investigators found that metabolic costs decreased the 

most with the RHEO KNEE followed by the C-Leg. 

Furthermore, the microprocessor-controlled knees 

demonstrated biomechanical advantages in several 

kinetic and kinematic parameters.38 EMG analysis 

revealed that there was decreased muscular activity of the gluteus medius on the affected side with 

the RHEO KNEE compared to the Mauch SNS and 

C-Leg. Finally, accelerometer data demonstrated 

smoother transitions from swing to stance and vice 

versa with the microprocessor knees.38 Although 

these studies seem to demonstrate differences 
548

Care of the Combat Amputee

among the knees studied, these results need further 

corroboration. This is highlighted by the current 

self-selection of conventional hydraulic knees by 

military amputees.

In regards to the prosthetic socket, only one C-level 

study was found that suggested increased symmetry of gait with vacuum versus suction sockets. The 

methodology of this study, however, was poor with no 

control for other prosthetic components. The literature 

analysis dealing with prosthetic mass supported the 

points already discussed in the section above.60

In summary, when it comes to prosthetic prescription guidelines, the clinician must rely on a consensus 

team decision and patient input due to the current 

paucity of scientific evidence for the prescription of 

prosthetic devices to injured service members.

Effect of Gait Training

A review of the Ovid database over the past 10 

years revealed only two articles on this topic, both by 

the same investigators.62,63 Both papers originate from 

the same research with the first paper describing the 

effect of a gait reeducation program on self-selected 

walking speed and pain-free or adaptive equipment 

free ambulation. The second paper described the effects of the gait reeducation program on sagittal plane 

kinetic and kinematic variables. 

The sample size for both papers was 9 transfemoral 

trauma or tumor-related amputees with a mean age of 

33 years. All had worn their respective prosthesis for 18 

months. Measurements were made at baseline (before 

the start of training), at the conclusion of the training program, and at 6 months follow-up. The training program 

averaged 10 months and incorporated traditional motor 

skills based gait training with a psychologically based 

self-awareness and confidence building outreach component. Training sessions occurred weekly for 1 to 2 hours. 

The investigators reported that the gait re-education 

program led to improvements in SSWV, gait symmetry, 

and an associated increase in the amount of muscle 

work on the amputated side. Three subjects were able 

to walk without walking aids, seven subjects learned 

to jog, and all had either a decrease or a disappearance 

of pre-training low back pain. These findings persisted 

until the 6-month follow-up. Although the findings 

described in these two studies are promising, more 

studies are needed in this area.

GAIT LABORATORY DEPENDENT THERAPEUTIC INTERVENTIONS 

Although rarely implemented in the past, the use 

of advanced technologies holds great promise for rehabilitation of the military amputee.64,65 The combined 

resources of real-time feedback and virtual reality are 

being implemented with the goal of improving gait 

performance. Current efforts are based on promising 

work that suggests that improved symmetry of loading 

and decreased metabolic cost can be attained in a relatively short period of time using visual feedback.64,65

It is anticipated that the use of the virtual reality and 

a real-time feedback could be effective in accelerating 

the rehabilitation process, especially with gait training. 

This method also could prove valuable in prosthetic 

device modifications, gait strategies used by amputees, 

and selection of an environment to achieve optimal 

efficient prosthetic gait. 

SUMMARY

This chapter has attempted to provide the reader 

with an overview of the terminology of gait, as well 

as observational and computerized gait assessment. 

The able-bodied “normal” gait has been discussed as 

a construct from which to understand gait deviations 

in the amputee. An indepth discussion of transtibial 

and transfemoral amputee gait has been provided with 

specific discussions concerning kinematic and kinetic 

parameters of joints throughout the gait cycle. Further 

discussion of specific considerations in amputee gait 

such as energy expenditure and prosthetic mass effect 

has also been provided. The chapter concludes with a 

discussion of how gait laboratory evaluation assists in 

the management of lower extremity amputees.

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