Updated 08/
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Over the past decade, technology and research have greatly expanded the functionality and aesthetics of prosthetic feet. Today, amputees have a wide array of feet from which to choose. Various models are designed for activities ranging from walking, dancing and running to cycling, golfing, swimming and even snow skiing. Heavier wood and steel materials have been replaced over the years by lightweight plastics, metal alloys and carbon-fiber composites. Much like the human foot, many of today’s prosthetic feet can store and return some of the energy generated during walking. Other key attributes included toe and heel springs that allow more natural movement at the ankle, shock absorption, multi-axial rotation, adjustable heel heights, and waterproof materials.
A number of factors must be considered when selecting the right foot/feet for your lifestyle. These factors include your amputation level, age, weight, foot size, activity level, goals and occupational needs.
Structurally, prosthetic feet can be divided into two groups: those with a rigid connection to the prosthetic shank (non-articulated) and those with a hinged ankle mechanism (articulated). In terms of function, prosthetic feet can be categorized into the following groups:
Although not all are discussed in this Fact Sheet, the following are definitions of terms you may hear when discussing various types of prostheses, fitting needs and activity requirements with your prosthetist and physician. This knowledge may help you choose which type of prosthesis is the most appropriate for you and your daily activities and needs. Never hesitate to ask for clarification from your prosthetist or physician if you do not understand something they say. You are an important part of your medical team.
Internal and External Rotation: Internal rotation refers to movement of a joint or body part toward the center of the body, while external rotation refers to the opposite rotation of a joint away from the body.
Dorsiflexion and Plantarflexion: The upward (dorsi) and downward (plantar) movements of the ankle and toes. These movements alternately enable the leg to move forward over the foot, pushing the forefoot to the ground as one takes a step.
Inversion and Eversion: The inward and outward, or side-to-side, motions of the ankle.
The most basic prosthetic feet come in two types: Solid Ankle Cushioned Heel (SACH) and Elastic Keel configurations. These designs consist of crepe neoprene or urethane foam molded over an inner keel and shaped to resemble a human foot. Because they have no hinged parts, these basic feet are relatively inexpensive, durable and virtually maintenance-free. These feet offer cushioning and energy absorption but do not store and return the same amount of energy as dynamic-response feet. SACH and elastic keel feet are generally prescribed for amputees who do a limited amount of walking with little variation in speed.
SACH Foot: The SACH is the simplest type of non-articulated foot. The name refers to a somewhat soft rubber heel wedge that mimics ankle action by compressing under load during the early part of the stance phase of walking. The keel is rigid, which provides midstance stability but little lateral movement. The SACH foot is available in various heel heights to match individual shoes with different heel heights.
Elastic (flexible) Keel Foot: This prosthetic foot allows motion similar to that of SACH feet. In addition, the forefoot is able to conform to uneven terrain but remains supportive and stable during standing and walking.
Articulated prosthetic feet may be single-axis or multi-axis in their design. “Axis” refers to motion in one or more of three different planes, similar to the movement of the natural foot. Prosthetic feet that have movement in two or three axes provide increased mobility at the ankle, which helps stabilize the user while navigating on uneven surfaces.
Single-Axis Foot: The articulated single axis foot contains an ankle joint that allows the foot to move up and down, enhancing knee stability. The more quickly the full sole of the foot is in contact with the ground, the more stable the prosthesis becomes. This is beneficial for users with higher levels of amputation (an amputation anywhere between the knee and hip). The wearer must actively control the prosthesis to prevent the knee from buckling, and the single-axis ankle/foot mechanism reduces the effort required to do so. Unfortunately, the single-axis ankle adds weight to the prosthesis, requires periodic servicing, and is slightly more expensive than the more basic SACH foot. A single-axis foot may be more appropriate for individuals where stability is a concern.
Multi-Axis Foot: Although similar to the single-axis foot in terms of weight, durability and cost, the multi-axis foot conforms better to uneven surfaces. In addition to the up and down mobility of the single-axis foot, a multi-axis foot can also move from side to side. Since the added ankle motion absorbs some of the stresses of walking, this helps protect both the skin and the prosthesis from wear and tear.
People with more active lifestyles typically prefer a more responsive foot. A dynamic-response foot is ideal for those individuals who can vary walking speed, change directions quickly or walk long distances. Dynamic-response feet store and release energy during the walking cycle by absorbing energy in the keel during the “roll-over” phase and then springing back to provide a subjective sense of push-off for the wearer. Additionally, they provide a more normal range of motion and a more symmetric gait. Some dynamic-response feet feature a split-toe design that further increases stability by mimicking the inversion/eversion movements of the human ankle and foot.
The comfort and responsiveness of a dynamic-response foot can also encourage an individual to advance from a more moderate activity level to a higher activity level, given the more natural feel of walking with this type of prosthetic foot. Further, some dynamic-response feet have been shown to reduce impact forces and stress upon the sound side foot and leg.
Microprocessor-controlled (MPC) feet are a fairly new category of prosthetic components. These foot/ankle components have small computer-controlled sensors that process information from both the individual’s limb and the surrounding environment to adjust to various needs. Based on information from input signals, these processors apply an algorithm, or set of rules, to make decisions about how the ankle or foot should respond in any given situation. The microprocessor provides instructions to various parts of the prosthesis in order to produce the desired function of the foot. Current MPC ankles use a variety of sensors, including ankle angle sensors, accelerometers, gyroscopes and torque sensors. The microprocessors in these systems then take the input signals and make decisions as to how to position the ankle, how to set the damping resistance in the ankle, and how to drive an ankle motor during stance phase (1).
The largest potential benefit of an MPC ankle/foot system over other prosthetic feet is the enhanced ability to react to varying environmental situations by providing different mechanical properties or alignments to improve the user’s balance and mobility. For example, non-MPC prosthetic feet work nicely on smooth, level terrain; however, they have a more limited ability to alter their mechanical properties or alignment when walking on slopes or other uneven surfaces. Powered feet provide propulsion during ambulation to enhance walking capabilities in real-time. Some specific models include software as well as options for connectivity to mobile devices through smart or computer apps. This allows the prosthetist and user to match the performance of the ankle/foot to various activities, allow for adjustments to the input gains and timing, and turn on or off certain features. All of these functions provide a more individualized experience by the user.
The ultimate goal of this class of prosthetic feet is to mimic the functions of the human foot. However, devices differ in their ability to accommodate for all environments and thus to the extent in which that accommodation can be achieved (2). Although these types of feet can coordinate the movements of the foot and ankle automatically, they do not directly communicate with the body. Microprocessor or powered prosthetic feet require batteries to power the chip, sensors, motors and actuators. Additionally, electronic parts associated with microprocessor systems make them more delicate than their passive counterparts. Many should not be used in water or in highly dusty or dirty environments. Due to the extra parts required by the addition of the microprocessor, they often weigh more than other prosthetic feet. Users may notice the mechanical clicks and sounds coming from the prosthesis as the microprocessor extrapolates information and adjusts various aspects of the ankle or foot. Finally, the higher level of technology and more intricate design of this class of prosthetic feet mean they may likely be the more expensive options on the market.
Just as there is no single tool perfectly suited for every job, there is no single foot that is perfect for every amputee. Knowing the available options will enable you to discuss this issue clearly with your prosthetist. Evaluate the pros and cons of different feet so you can make the best choice for your individual aspirations and abilities. In comparing the potential benefits of microprocessor-controlled systems over other systems, physicians and prosthetists should focus on the functional aspects of the prosthetic foot and its level of appropriateness, given the user’s individualized needs and goals.
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The gait patterns of unilateral below-knee amputees wearing prostheses with either a SACH foot or a single axis foot were compared. A temporary below-knee prosthesis was fabricated for each subject using plaster of Paris and Plastazote for the socket, a pylon and an artificial foot. Eight subjects were filmed at two separate sessions, one in which the SACH foot was worn on their prosthesis and one with the single axis foot on their prosthesis.
Measurements of the normal leg with a SACH foot on the prosthetic limb were compared to measurements of the normal leg with a single axis foot on the prosthesis. Measurements of the prosthetic leg with both devices were also compared. A one tailed t test (p<.05) was used to determine statistical significance of the results obtained in six measurements of lower limb joint angles and on the percentage of the time of gait cycle for stance and swing phase of the prosthetic leg.
Discussion centres on the interpretation of the results from both statistical and clinical points of view. Major differences (excepting the ankle at foot-flat) between the prosthetic devices were not found.
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The two artificial feet used in this study were the single axis foot and the SACH (solid ankle cushion heel) foot. These devices seem to be the most commonly used at present. Throughout the 's the SACH foot was the artificial foot of choice in North America when fitting the majority of below-knee amputees (Fishman, et al. ). However, the single axis foot, which was first used during World War 1, appeared to be capable of more closely simulating human gait, due to the dorsiflexion and plantarflexion movement of the device.
The choice of prosthetic foot is important to the amputee and to the amputee clinic team prescribing the prosthesis. According to Statistics Canada, there were 1,198 below-knee amputations performed in Canada in . Patients in the 55 to 67 age group are usually amputated for peripheral vascular disease which accounts for 80 per cent of the total number of amputations (Hunter and Waddell, ).
Breakey () observed the effect of lost ankle movement on the gait of below-knee amputees by comparing the gait of five normal subjects to the gait of five unilateral below-knee amputees. These amputees were wearing a patellar tendon bearing prosthesis with a SACH foot. Breakey () suggested that the lost ankle movement affects the knee movement and foot timing. He stated that knee flexion was decreased to 7 degrees from a normal value of 17 degrees during the stance phase, decreased to 30 degrees from a normal value of 37 to 43 degrees during toe-off and decreased to 57 degrees from a normal value of 61 to 68 degrees during the swing phase of gait.
Stance phase occurred for 57% of the time of the gait cycle for the involved limb and 63% for the uninvolved limb in Breakey's amputee subjects.
Robinson et al. () tested 19 unilateral below-knee amputees, mean age 43, wearing a SACH foot on their patellar tendon bearing prosthesis. They observed a mean stride length of 1.32 metres, a mean step length from uninvolved to involved limb of 0.68 metres, and a mean step length from involved to uninvolved of 0.63 metres. An increased amount of time spent on the uninvolved limb compared to the involved limb contributes to an increase in stance phase. The mean walking velocity of these subjects was 1.07 m/s.
Eight males were selected according to the following criteria: unilateral below-knee amputees wearing a patellar tendon bearing prosthesis with cuff suspension. These subjects were between the ages of 55 and 67, in good general health and had no skin problems with their stump.
Each amputee was fitted with a temporary prosthesis to accommodate the interchange of prosthetic feet. The temporary prosthesis consisted of a plaster socket lined with Plastazote, cuff suspension, a pylon and prosthetic foot. The socket was fabricated and aligned on the pylon and foot following the same principles as a permanent patellar tendon bearing prosthesis.
The selection of the first prosthetic foot to be measured was made according to the availability of the prosthetic feet. The subjects were tested two days after fitting of a prosthetic foot if the foot was the same design as the one on their permanent prosthesis. A time lapse of one week was allowed if the prosthetic foot was not the same design as the one worn on their permanent prosthesis. The first prosthetic foot was changed after the filming was completed. The second foot was aligned on the prosthesis and the subject given a date for the second filming.
A distance of approximately six metres (on level ground) was used as a walking zone allowing the subjects to complete three gait cycles. The subjects were filmed simultaneously from lateral and frontal perspectives. The anterior view (Bolex camera, 64FPS) was used to check knee angle measurements (varus, valgus), lateral deviation of the trunk and shoulder elevation. The subjects were filmed twice with the normal (non-amputated leg) closest to the side camera (Locam camera, 86FPS) and twice with the prosthetic leg closest to the camera.
A one foot measurement board was filmed in the centre of the walkway and used to convert digitized co-ordinates. The processed film was projected on to a digitizing tablet by a stop action projector. The nineteen body parts necessary for the computer to calculate the centre of mass were digitized and recorded on paper tape for each subject at heel-strike and every fifth frame following until the gait cycle for the leg was completed. The nineteen body parts in sequence, were the head, sternum, crotch, right shoulder, elbow, wrist and hand, left shoulder, elbow, wrist and hand, right hip, knee, ankle and foot and left hip, knee, ankle and foot. The paper tape was fed into a computer terminal and data points stored on a magnetic disc. Key punch cards were used to obtain the program output from the computer which included the path of the centre of mass for each subject.
The measurements obtained in this study to compare the gait of the amputee subjects were as follows:
The vertical displacement and velocity of the centre of mass of each subject on each trial were obtained from the computer printout.
The lower limb joint angles were measured by projecting the lateral view films on to a wall using a stop action projector. The selected angles of the hip, knee and ankle were measured using a goniometer. The joint angles for heel-strike were measured as the heel came in contact with the floor following swing phase. Foot-flat was measured as soon as the entire foot came in contact with the floor. As the hip moved directly over the foot, angle measurements for mid-stance were taken. The angle measurements for heel-off, as the heel left the ground and toe-off, as the toe left the ground completed the stance phase measurements.
During swing phase the angle measurements were taken as the knee passed directly under the hip for acceleration, as the foot passed directly under the hip for mid-swing and as the knee ceased to extend for deceleration.
The percentage of time of gait cycle of stance phase, swing phase and double support phase and the time of the gait cycle were calculated using the number of fames and the frame rate of the film.
Step length and stride length were obtained using the digitizing tablet and the projected film.
The measurements obtained when the subjects were wearing a single axis foot on their prosthesis were compared to those obtained when they were wearing a SACH foot on their prosthesis. A paired t-test was used to determine whether the two means were significantly different. A probability level of 0.5 was selected for a 1-tailed t test. The degrees of freedom for all comparisons was seven. Statistical significance was obtained in six comparisons of lower limb joint angles and the percentage of the time of gait cycle for stance and swing phase of the prosthetic leg (Table 1). The mean velocity of the body's centre of mass for both the SACH foot and the single axis foot measurements was 1.22 m/s. The difference in velocity comparing the subjects SACH foot gait and single axis gait was no greater than 0.2 m/s. (Table 2).
Although statistical significance was obtained in six comparisons, only one variable was felt to have any clinical significance. The determination of clinical significance was based on research by Murray () on 30 subjects which identified one standard deviation of ankle angle measurements to be approximately 7°, knee angle to be 5° and hip angle to be 11° at an average walking speed of 1.39 m/s. In this study the ankle angle at foot-flat of the prosthetic leg measurements showed a difference of 6.5° between the single axis foot and the SACH foot measurements. The design of the single axis foot permits a greater range of plantarflexion and dorsiflexion around a transverse ankle axis with the range limited by posterior and anterior rubber bumpers. The SACH foot permits a limited range of plantarflexion through the compression of the posterior rubber heel insert. The remaining statistically significant lower limb joint angles were not considered clinically significant since a measurement difference of 4° or less was obtained.
The percentage of time of gait cycle for stance and swing phase of the prosthetic leg fell within the reported normal ranges of 60 and 40 per cent (Drillis, ) and 62 and 38 per cent (Peizer et al. ) for stance and swing phases. Although the difference obtained was statistically significant, once again it was not felt to be of clinical significance.
In this study, interchanging the prosthetic foot on a prosthesis did not appear to have a significant effect on the gait patterns of unilateral below-knee amputees. The ankle of the prosthetic foot during the foot flat phase of gait showed a significant statistical and clinical difference.
The authors are grateful to the Nova Scotia Branch, War Amputations of Canada for their support and in providing funds for this study.
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