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Section 1: General Topics
Figure 3
across the gait cycle (0% to 100%) during walking. The black solid line represents group mean data for normal, intact individuals without limb loss.
The black dashed line represents group mean data for a passive (hydraulic) prosthetic knee. The red dashed line represents group mean data for a
microprocessor-controlled prosthetic knee. The black dash-dot line represents group mean data for a passive prosthetic foot-ankle. The red dashdot-dot line represents group mean data for another passive prosthetic foot-ankle. The green dotted line represents group mean data for a powered
prosthetic foot-ankle. Conceptually, moments are eort and power is success.
normal intact ankle during gait (Figure
3, I). This remains a critical problem for
those with limb loss because the pushoff power from the prosthetic side does
not accelerate the body sufficiently to
counteract the deceleration caused by
the leading native limb at foot contact.
This results in a change of momentum
that is costly in both mechanical and
metabolic terms for those with limb loss.
A more effective prosthetic foot will also
influence cosmesis, with the appearance
of a more normal gait pattern and a less
pronounced limp.
Atlas of Amputations and Limb Deciencies, Fourth Edition © 2016 American Academy of Orthopaedic Surgeons
72
Graphs illustrating sagittal plane motions; moments; and power of the hip (A, D, and G), knee (B, E, and H) and ankle (C, F, and I)
The sagittal motion of the ankle has
been described as consisting of three
“rocker” phases: the first rocker, from
initial contact on the heel to foot flat
with ankle plantar flexion; the second
rocker, with the tibia advancing over
the plantigrade foot in progressive
dorsiflexion; and the third rocker, in
which the ankle plantarflexes in late
50,51
stance.
The prosthetic foot-ankle
system can emulate the first and second
ankle rockers quite closely, but it does
not actively plantarflex as much as the
intact biologic ankle during the third
rocker. No prosthetic foot plantarflexes
past neutral; the intact biologic ankle
plantarflexes to 30° to 40° during the
end of the stance phase of walking. Efficient amputee gait depends on the user’s
ability to intuitively use the spring characteristics of the prosthetic foot to mimic
the push-off forces of the biologic ankle
observed during the late stance phase.
The sagittal moment profile of the
foot during walking begins at near zero
or slightly negative, indicating contact
with the heel and a brief dorsiflexion
moment. The moment crosses the axis
Chapter 6: Amputee Gait: Normal and Abnormal
and moves into a plantar flexion moment as foot flat is achieved and the GRF
vector passes directly through the ankle
joint center. For an instant during midstance, all forces at the ankle are compressive, but as stance progresses, the
sagittal moment becomes more positive,
indicating substantial effort by the plantar flexors to restrain the shank segment
and guide the ankle into controlled dorsiflexion. This might be considered an
eccentric contraction of the triceps surae
to restrain the forward progression of
the tibia over the foot, but recent ultrasound studies suggest that the fascicles
of the gastrocnemius during walking
are actually isometric and that most of
the observed ankle dorsiflexion in midstance and terminal stance are the result
of tendon stretch.52 However, a weak triceps surae can result in a calcaneal gait
or a crouch gait pattern. Compensation
can be attempted by using the knee extensors. As stance progresses, enough
force is eventually generated to lift the
heel and initiate plantar flexion in late
stance and preswing. For the intact ankle, small amounts of joint power are absorbed as stance progresses, indicating
an eccentric contraction of the plantar
flexors, but in late stance and preswing,
a substantial power generation peak occurs, indicating a concentric contraction
by the plantar flexors to provide pushoff energy and propel the body forward
into the swing phase.
The subtalar and midtarsal joints
that control the coronal plane motions
of the foot relative to the ankle in the
normal biologic lower limb have never
been emulated with biomimetic accuracy in prosthetic feet. Some prosthetic
feet are less stiff in the coronal plane
and can passively accommodate side
slopes or leaning while turning. For
some individuals, however, these feet
feel somewhat unstable in the coronal
plane. Personalized prosthetic prescriptions that consider the range of designs,
user activities, and user functional
performance goals are recommended.
Some individuals (especially children
with limb loss) will choose a carbon
fiber running blade as their prosthetic
foot. This type of prosthesis is inherently unstable in stance, but can bend
and flex to a large degree and provide
much higher push-off power because
it has no cosmetic foot cover or cushioned shoe. Younger children especially
seem to value the trade-off, accepting
instability while standing and enjoying
the ability to run as fast as their peers
during outdoor play. Running-specific
prosthetic feet are not covered by the
Centers for Medicare and Medicaid Services or private health insurers, but may
be an appropriate prescription for some
highly active prosthetic users, usually as
a completely separate socket-pylon-foot
system that can be used for running.
Infants with limb loss who lack a knee
have historically not been provided an
articulating prosthetic knee until they
begin to walk. However, recent research
suggests that infants with transfemoral
limb loss appear to benefit from a prosthetic knee that can bend to accommodate crawling.53 Such a knee also appears
to improve mobility during play, as well
as the achievement of the developmental
milestones of pull-to-stand and walking
in infants with limb loss.
The power generated in the sagittal
plane by a prosthetic ankle (calculated from computerized gait analysis)
during late stance phase is usually less
than 50% of that of the intact biologic
ankle during walking,
3,18,21, 22,32,54
with
no net power generation (only powered and experimental passive spring
prosthetic ankles can achieve normal
power generation). The one exception
to limited push-off power is the new,
powered prosthetic ankle that can provide plantar flexion power during walk-
54,55
ing.
Powered prosthetic ankles have
tremendous potential, but currently
are heavy, noisy, have limited battery
life, are not reimbursed by the Centers
for Medicare and Medicaid Services or
private health insurers, and require the
user to successfully use increased pushoff energy.54 A successful prototype of an
MPC prosthetic foot uses the collision
energy from initial contact and recycles
this into propulsion energy during late
3,56,57
stance.
A few studies have shown
improvements in ankle power generation with these new MPC or powered
prosthetic ankles,
costs of walking appear unchanged,
3,54 -57
but the metabolic
3,54
suggesting that gait efficiency remains
elevated for prosthetic users irrespective
of the sophistication of the prosthetic
component.
Prosthetic Foot Type
and Ankle Kinematics
and Kinetics
Prosthetic components are provided
based on established criteria related
to the patient’s functional level. The
Medicare/Medicaid Functional Classification Level58 (MFCL K-level) is rated
on a scale from 0 to 4. The lower functioning levels (K0, K1, and K2) receive
less advanced, less expensive prosthetic
components (for example, solid ankle–
cushioned heel [SACH] feet or passive
hydraulic knees). The higher functioning levels (K3 and K4) receive technologically advanced prosthetic components
(for example, energy storage and return
feet or MPC elements). The physician
establishes the individual’s K-level using
observational gait analysis and patient
input. Although ambulation attributes
have been formally described,58 no
reliable or valid method exists to precisely determine K-level in a patient
with limb loss. The Amputee Mobility
Predictor (AMP)59 has some published
validity, but the evidence was based
on clinician ratings that have substantial errors. The 6-minute walk test60 is
highly correlated with the AMP, takes
less time to administer than the AMP,
and is perhaps the most objective and
least time- consuming functional level
assessment available. A version of the
AMP that is used for individuals with recent amputations (AMPnoPRO)59 can be
© 2016 American Academy of Orthopaedic Surgeons Atlas of Amputations and Limb Deciencies, Fourth Edition
73
Section 1: General Topics
quite valuable in predicting successful
functional level for an individual who
has never used a prosthetic limb.
The functional level of the individual
with limb loss is used to determine the
type of prosthetic components that can
be provided, and these components can
ultimately influence the joint kinematics
and kinetics of walking. Passive prosthetic feet with advanced carbon fiber
keels can flex and recoil substantially
during late stance and preswing, but
older SACH designs have wood or polyethylene keels with foam foot shells that
compress during the late stance phase.
SACH feet have a much lower power
generation peak in late stance than carbon fiber prosthetic foot designs,
21,22
and
should be prescribed for those with lower functional levels (K1 and K2). The
degree of forefoot keel flexibility can
be engineered to match the weight and
walking speed of the individual, but
rigorous data are still needed to inform
the prescribing physician about the appropriate forefoot stiffness of specific
models of prosthetic feet. The prescription of prosthetic feet is usually based on
professional clinical experience, without
using objective data that can be matched
with a specific patient profile. A team
approach with input from the user, the
prosthetist, and the prescribing physician is likely the most effective model of
care in choosing the most cost-effective
and mobility-empowering prosthetic
components.
Knee Sagittal Kinematics
Normal Knee Motion
The sagittal kinematics of the intact biologic knee over the gait cycle begins with
the knee in slight flexion at initial contact, flexes to approximately 15° (more at
faster walking speeds) at approximately
25% of the gait cycle, and extends again
to almost full extension late in the stance
phase before beginning to flex again
during preswing (Figure 3, B). At footoff, the knee is in approximately 30° of
flexion, continues to approximately 55°
in midswing, and extends until reaching
full extension late in swing before initiating slight flexion just before initial
ground contact.
Transtibial Limb Loss
(Retained Biologic Knee)
For those with transtibial limb loss, the
preserved biologic knee typically replicates the sagittal plane motion of the
intact biologic knee during walking,
including a consistent, recognizable
flexion wave during early stance. Most
transtibial prosthetic users have intact
knee extensors (quadriceps) and knee
flexors (hamstrings) and normal motor
function. The choice of prosthetic foot
can influence the amount of extensor
or flexor moments applied to the knee
during walking, but most transtibial
prosthetic users can manage the small
discrepancies in knee motion that are induced by prosthetic feet. Most transtibial
prosthetic users are able to walk with
knee kinematics on their involved side
similar to that of an intact biologic knee.
If the normal functional performance
of the biologic knee is compromised in
some way in those with transtibial limb
loss (for example, weakness, joint deformity, or motor nerve injury), the forefoot
stiffness of the prosthetic foot might be
increased to provide greater extensor
moments in midstance, thus improving
knee extension rather than a flex and
recoil pattern to provide push-off power.
This compromise between support and
propulsion must be carefully considered
for each individual.
The residual limb length of any transtibial amputation should be optimized,
if possible, to make community ambulation more functional. If the amputation results in a residual limb that is too
short, control of the prosthesis during
walking will require very high pressures
at the residual limb–socket interface
that will increase the risk of blisters,
sores, and ulcers. If the residual limb
is too long, the choice of prosthetic feet
for the individual will be limited and
require difficult design and fabrication
work by the prosthetist. Boyd amputations or Syme ankle disarticulations that
allow distal weight bearing are generally
advantageous, but if they are excessively
long they can limit prosthetic foot prescription options that provide the greatest flex and recoil during walking. Seven
to 10 inches of clearance between the
end of the residual limb and the floor
allow the greatest range of possible prosthetic components to meet the prosthetic
user’s range of desired functional performance goals. These decisions regarding
the amputation site are often driven by
medical or biomechanical factors, and
residual limb length optimization may
not be possible for every patient. However, the effects of residual limb length
on pressure-to- moment relationships
during ambulation and on prosthetic
component choices should be considered by the treating surgeon.
The choice of shoes by the prosthetic
user is beyond the control of healthcare
providers, but it can radically alter the
kinematics and kinetics of the prosthetic limb and the compensatory mechanisms of the retained biologic knee.
A shoe with a higher heel will tend to
incline the prosthetic limb forward and
require additional knee extensor muscle
activity to maintain appropriate knee
extension. A few MPC feet currently on
the market can accommodate different
heel heights and can be set by the prosthetic user whenever he or she changes
shoes.
Transfemoral Limb Loss
(Absent Biologic Knee)
Individuals with transfemoral limb loss
must use a prosthetic knee joint to emulate the sagittal kinematics needed for
walking. The prosthetic knees of individuals with limb loss from transfemoral amputation or knee disarticulation
usually lack sagittal plane flexion during
early stance (Figure 3, B). The limited
flexion during early stance occurs because most prosthetic knees are unable
Atlas of Amputations and Limb Deciencies, Fourth Edition © 2016 American Academy of Orthopaedic Surgeons
74
Chapter 6: Amputee Gait: Normal and Abnormal
to flex eccentrically during the early
portions of stance phase. Even users of
MPC knees, which are designed to have
this knee flexion wave in early stance,
do not often exhibit a normal kinematic
knee pattern.
61-63
The inability to walk
with a normal kinematic knee pattern
can be the result of a lack of trust in the
knee by the user or an issue with knee
alignment in the sagittal plane, or it can
possibly be related to software parameters set by the prosthetist. The ability to
achieve eccentric flexion in early stance
phase in prosthetic knees has not been
demonstrated convincingly.
24,61,64,65
MPC knees have demonstrated a
decrease in knee hyperflexion (flexion greater than 60°) during the swing
phase.61 This decrease makes achieving
full knee extension before initial contact more likely at a range of different
walking speeds, a key feature to avoid
knee buckling and falls in individuals
with transfemoral limb loss. Hydraulic
or mechanical prosthetic knees usually
exhibit knee hyperflexion in swing61 but
can be adjusted to increase joint dampening to normalize the sagittal kinematics of the prosthetic knee joint during
swing. However, knee flexion can only
be optimized for a specific flexion force
and, therefore, a specific walking speed;
the induced pendular mechanics of
the prosthetic knee at faster or slower
walking speed needs some dynamic
dampening to accomplish the correct
magnitude and timing of the peak flexion at these different speeds. For those
individuals with demonstrated multiple
walking speeds, an MPC knee can offer
true functional and safety advantages
because of this swing phase adaptation feature.61 Payers only reimburse an
MPC knee if the patient is at level K3 or
higher, uses multiple walking speeds,
and can competently handle most community environmental barriers, such as
curbs, ramps, and stairs. Some individu
als with lower functional levels (K2) can
achieve multiple walking speeds (K3)
when provided with an MPC knee.
24,62
In addition, there is some evidence that
individuals at the K2 functional level
can benefit from an MPC prosthetic
knee that can dampen flexion moments
to maintain stability during stumble re
covery and possibly improve the safety
of walking on ramps.
1
The shoe choice for a transfemoral
prosthesis user has a more dramatic effect on the kinematics and kinetics of
walking than shoe choice for a transtibial prosthesis user. Shoes with increased heel height will tend to cause the
prosthetic knee to buckle and collapse
in an individual with a transfemoral
amputation, increasing the risk of falls.
Therefore, the range of possible shoe
choices is more limited in transfemoral
prosthesis users, although some MPC
prosthetic feet permit the transfemoral
amputee to adjust the neutral position of
the ankle to accommodate different heel
heights, allowing for a greater range of
footwear choices. This is an important
cosmetic and social inclusion issue for
formal occasions (high-heeled shoes) or
outdoor work activities (boots for manual labor) for prosthesis users.
Normal Knee Kinetics
The kinetic patterns of the knee have
complex waveforms during walking but
primarily provide support for the body
during stance. Early stance often has a
flexion moment even though the knee
is near full extension, suggesting some
stress on the posterior capsule because
the knee must resist further extension
(Figure 3, E). The sagittal knee moment
quickly crosses the axis into an extension moment, and substantial quadriceps activity is needed to maintain
upright stance despite external forces
that would cause the knee to collapse.
Prosthetic Knee Kinetics
When an individual with a prosthetic
-
knee joint walks, the initial knee flexion
moment generally locks the prosthetic
knee in extension and it remains there
throughout most of the stance phase.
This corresponds to the obvious lack
of knee flexion during early stance observed in the sagittal knee motion of
prosthetic knee users, even with MPC
-
knees designed to have knee flexion
in early stance. As stance progresses,
the flexion moment increases beyond
the magnitude seen in the intact native
knee, indicating substantial resistance
to hyperextension by the prosthetic
knee (Figure 3, E). In late stance, both
normal and prosthetic knee moments
have a brief extension moment just as
knee flexion is initiated.
Hip Sagittal Kinematics
Sagittal hip motion is primarily sinusoidal during walking. Stance begins
with approximately 40° of flexion at
initial contact, with the hip extending
to neutral at midstance and moving into
approximately 5° to 20° of extension
during terminal stance, with greater
hip extension at longer stride lengths
(faster walking speeds). The hip begins
to flex before foot-off and moves through
neutral during midswing until the hip
reaches full flexion of approximately 40°
during the late swing phase.
An individual who lacks a knee
joint must use the femoral residual
limb to apply pressure to the interior
of the prosthetic socket to simultaneously achieve the sinusoidal hip motion
during gait and control the prosthetic knee joint. Because the hamstrings
no longer function as hip extensors
for most transfemoral amputees, the
gluteal muscles must provide the hip
extensor forces during the first half of
stance and the iliopsoas must provide
the hip flexor forces during late stance
and swing. Sagittal hip motion in individuals with transfemoral limb loss is
often shifted toward flexion and has a
somewhat reduced arc of sagittal motion, primarily because of a lack of full
extension in late stance (Figure 3, A).
The lack of push-off power from the
prosthetic foot in late stance results in
lower energy transferred to the limb to
© 2016 American Academy of Orthopaedic Surgeons Atlas of Amputations and Limb Deciencies, Fourth Edition
75
Section 1: General Topics
initiate swing in the transfemoral prosthesis user. Greatly increased hip flexor
muscle activity can improve the range
of motion of the hip in individuals with
transfemoral limb loss, but generally the
increased pressures in the distal anterior region of the residual limb–socket
interface are uncomfortable or perhaps
injurious to the tissue over time. The
three basic mechanisms to advance the
limb into swing during walking are ankle push-off energy transferred to the
tibial and femoral segments (distal to
proximal strategy), hip flexor muscle
activity to accelerate the thigh into
flexion and initiate swing (proximal to
distal strategy), and deceleration of the
pelvis and trunk segments to transfer
momentum to the training limb on contact with the lead limb (proximal to distal strategy). Individuals without lower
limb loss can manage all three methods
with some form of energy and control
optimization,66 but the range of step-tostep variance and acceptable solutions
to the balance between the forces that
initiate swing is not completely understood. Of all the tasks in human gait,
swing initiation is likely less important
than managing the step-to-step transitions during double support without
wasting mechanical energy and therefore increasing metabolic cost.
46
An individual with a hip disarticulation uses pelvic tilt in the sagittal
plane to facilitate the motion of the
missing thigh segment during ambulation and must learn a complex set of
trunk maneuvers to maintain forward
progress during walking. As a consequence, low back pain is a frequent
symptom in individuals with more
proximal limb loss.
67-71
Individuals with
a short femoral residual limb segment
produce hip and pelvic motions during
walking that are similar to those of individuals with a hip disarticulation;
this makes successful community ambulation more difficult. Every surgical
and medical effort should be made to
maintain an adequate femoral residual
limb segment after limb loss. Without
adequate length, the pressures placed
on the tissues of the residual limb are
too high to be tolerated during longer
periods of walking. The hip extensors
and flexors are at an extreme mechanical disadvantage if the femoral residual limb is too short. Many individuals
with hip disarticulation or a short femoral residual limb will need alternative
mobility solutions, such as a scooter,
wheelchair, or Lofstrand crutches for
some portion of their daily activities;
this is especially true as an individual
ages.
Hip Sagittal Kinetics
Hip kinetics in individuals with limb
loss at the knee and higher levels reflect
the increased effort of the hip in moving the prosthetic knee and ankle joints.
The early stance hip extension moments
(gluteal muscle activation) in the intact
hip joint are similar to those in individuals with transfemoral limb loss. As
stance progresses, individuals with more
proximal limb loss must apply greater
hip flexion moments (iliopsoas muscle
activation) to initiate swing phase knee
flexion and limb advancement because
they lack the ankle push-off power
that can transfer energy to the swing
limb.72 The action can be taxing, with
peak sagittal power at the hip more than
double that of normal, intact gait. This
can contribute to increased pressure
at the residual limb–socket interface
near the distal anterior region, a location of frequent skin irritation for some
transfemoral amputees. Compensatory
actions to advance the prosthetic limb
into swing are often observed visually,
including “hip-hiking,” or Trendelen
burg gait with circumduction. (Trendelenburg gait or lateral trunk lean over
the stance limb is more often related to
reducing coronal moments about the hip
and medial- distal and proximal-lateral
socket pressure discomfort in individuals with transfemoral limb loss.)
Ground Reaction Forces
During the stance phase of gait, an interaction between the individual and the
ground results in force being applied
to the distal foot (Figure 4). The reaction force measured during computerized gait analysis is the GRF. The GRF
has three components that aid in forward motion during walking: anteriorposterior, medial-lateral, and vertical.
The largest force during walking is the
vertical GRF. The vertical GRF is a bimodal curve with each peak reaching
approximately 120% of body mass in
healthy walking. The first peak in the
vertical ground reaction curve measures
the weight acceptance of the limb, and
the second peak measures the propulsive force needed to move the body
forward. Any disturbance in the timing
or magnitude of these GRF peaks indicates a kinematic deviation from normal
walking patterns.
Individuals with limb loss generally
have similar GRF patterns, but close
examination reveals typical deviations.
The late stance vertical GRF peak is often lower in amputees than individuals
without limb loss, due in part to the lack
of active push-off power from almost all
prosthetic feet. Typically, slower walking speeds result in a less-pronounced
bimodal pattern for the vertical GRF in
individuals with limb loss. Wider foot
placement also results in somewhat
greater medial-lateral force pattern, but
the magnitude remains relatively low.
Activities of Daily
Living and Mobility
Data from computerized gait analyses
specifically focus on walking straight
-
ahead and can help in understanding a
prosthesis user’s ambulation challenges
while walking straight ahead. Sagittal
plane joint motions and forces provide
most of the motion and allow an individual to walk in a straight line. However,
observations of community ambulation
and function suggest that most functional human locomotion involves a few
Atlas of Amputations and Limb Deciencies, Fourth Edition © 2016 American Academy of Orthopaedic Surgeons
76
Chapter 6: Amputee Gait: Normal and Abnormal
Figure 4
the gait cycle (0%–100%) for a normal, intact individual without lower limb loss (black solid line)
and a prosthet ic device user with a transtib ial amputation (red dash- dot-dot line). Note the reduced
peak vertical force in the prosthetic device user ’s data during late stance, suggesting exing of the
forefoot keel of the prosthetic foot.
steps in a row before stopping or turn-
73,74
ing.
for daily living involves rising from a
seated posture, initiating gait, accelerating, turning, decelerating, terminating
gait, climbing stairs, maneuvering over
and around obstacles, and returning to a
standing or seated posture. This is likely
a function of western urban architecture, but 4 ± 1 steps is the most common
walking bout (number of steps before
stopping), accounting for approximately
15% of all bouts for a wide range of individuals with differing activity levels.
Both sedentary and active individuals show a substantial bias toward short
walking bouts; 40% of all walking bouts
consist of 12 ± 1 steps or fewer in a row,
and 75% of all walking bouts consist
of 40 ± 1 steps or fewer in a row.75 In
addition, short periods without ambu-
Line graph illustrating vertical ground reaction force during walking plotted across
only once per day but for some individ-
Much of the mobility required
uals can include 8% to 20% of the total
daily steps.75 Because of the typical challenges for everyday activities, stability
and maneuverability of the individual
with limb loss to walk short distances
with competence appears more crucial
than energy efficiency and gait economy, although the latter is still important. For children with limb loss, the
ability to initiate running suddenly
to keep up with friends appears more
crucial than stability. Given the choice,
75
many children opt for a running limb
as their primary prosthesis, even if it
results in more falls and makes standing still more difficult. Because most
development in children is play-based,
enabling high-energy movements seems
an important goal in prescribing pros-
thetic components.
lation occur frequently throughout the
day. Long-duration walking behavior
for exercise or long-distance transportation does not occur often, even for individuals who exercise regularly. These
long-duration exercise bouts may occur
Summary
Computerized gait analysis can pro-
vide comprehensive information on the
joint motions and forces during walk-
ing for individuals with limb loss, but
additional ambulatory challenges exist
within the community for prosthesis
users. A comprehensive evaluation of
mobility is recommended to maximize
functional status in prosthesis users
with lower limb loss. When determining
the effect of a prosthesis on gait mechanics, it is important to examine the effect
across the entire lower limb. The effect
of a prosthesis is not isolated to a single
joint; therefore, these alterations need
to be considered when determining the
optimal treatment plan for each patient.
Computerized three- dimensional gait
analysis can document the compensatory kinematics and kinetics in individuals with lower limb loss, and it can
inform clinical decision making when
prescribing prosthetic components and
evaluating the outcomes of surgical interventions. Computerized gait analysis is also the primary method used in
scientific research into the comparative
effectiveness of prosthetic component
treatment choices. Understanding this
literature can improve an orthopaedic
surgeon’s ability to provide clinical care
for individuals with limb loss and can be
essential in helping the surgeon understand the scientific basis for the development of novel prosthetic components.
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Atlas of Amputations and Limb Deciencies, Fourth Edition © 2016 American Academy of Orthopaedic Surgeons
80
Chapter 7
Clinical Considerations of Observational Gait Analysis
John T. Brinkmann, MA, CPO/L, FAAOP Phillip M. Stevens, MEd, CPO, FAAOP
Abstract
It is helpful to understand the clinical concepts related to the alignment of lower limb prostheses and observational gait assessment along with principles related to the orientation
of prosthetic components and the resulting gait patterns. To achieve acceptable alignment
and minimize gait deviations in patients with transtibial and transfemoral amputations,
the prosthetist can adopt a methodic approach to dynamic alignment that may include
structured assessment tools and incorporate video assessment.
Keywords: gait assessment; prosthetic gait deviations
Introduction
The overall goal of rehabilitation after
lower limb amputation is to support a
level of function commensurate with
the individual’s functional capabilities
before amputation. In terms of walking
ability, the goal is to match as closely
as possible the patient’s pre-amputation
kinetic and kinematic gait parameters.
However, even after successful rehabilitation, practitioners should expect that
inappropriate distribution of forces on
the residual limb, increased loading of
the contralateral limb, and increased
energy expenditure.
3-5
The long-term
altered biomechanics observed in the
gait of prosthesis users can contribute to the development of orthopaedic
problems, including osteoarthritis.
6,7
Gait deviations should be normalized
as much as possible to minimize these
deleterious effects.
structural asymmetries caused by the
amputation will result in kinetic and kinematic asymmetries.
1,2
Gait deviations
are observed and addressed by applying
established clinical principles and professional judgment to arrive at the most
appropriate gait pattern for each individual. A study guide is included at the
end of this chapter to identify common
gait deviations and causes.
Observational Gait Analysis
Because practitioners in most clinical
settings do not have access to threedimensional instrumented gait analysis, they rely on unassisted observation
as the primary method for identifying
pathologic gait patterns. Several limitations are inherent in observational
assessment. First, many factors that
influence a user’s gait are not directly
Consequences of
Gait Deviations
Prosthetic gait that deviates from normal parameters has detrimental physical and functional effects, including
Mr. Brinkmann is a faculty member at Northwestern University’s Prosthetic and Orthotic Center
and serves as a board member, owner, ocer, or committee member of the American Academy of
Orthotists and Prosthetists. Mr. Stevens is an employee of the Hanger Clinic and serves as a board
member, owne r, o cer, or committee member of the American Academy of O rthotists and P rosthetists.
observable. For example, the pressure
of the residual limb against the socket
and the corresponding joint moments
caused by those pressures can only be
observed using equipment and methods
not available in most clinical settings.
Second, observational gait assessment
has been demonstrated to have limited
validity and reliability for gait-related
events in neuromuscular and musculoskeletal pathologies.
8,9
Studies focused
specifically on patients with lower limb
prostheses have confirmed this limita-
10,11
tion.
However, as the prosthetic prac
titioner is currently limited by available
resources and techniques to achieve an
optimally fitting and functioning prosthesis, observational gait assessment
remains the primary strategy used by
clinicians when assessing the gait of
prosthesis users at the initial fitting and
subsequent appointments.
Prosthetic Alignment
Prosthetic components are assembled
in a particular spatial relationship to
each other, commonly referred to as the
alignment of the prosthesis. The alignment process typically involves three
distinct stages. During bench alignment,
components are assembled and attached
to the socket based on traditional protocols, the manufacturers’ recommendations, and the clinical judgment of
the prosthetist. During the static alignment stage, adjustments are performed
during weight bearing without ambulation, with a focus on socket fit and
gross alignment. Dynamic alignment
involves observation of the ambulating
patient and optimization of the spatial
relationship of the components.
One of the key elements of dynamic
alignment involves using observational
assessment and user feedback to detect
and reduce common patterns of pathologic gait (deviations) that indicate the
need to change the spatial orientation of
-
© 2016 American Academy of Orthopaedic Surgeons Atlas of Amputations and Limb Deciencies, Fourth Edition
81
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