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Terminology in Prosthetic Foot Design and Evaluation

Haideri, Nasreen F. PhD

JPO Journal of Prosthetics and Orthotics: October 2005 - Volume 17 - Issue 4 - p S12-S16

NASREEN F. HAIDERI, PhD, is affiliated with the Research Department, Texas Scottish Rite Hospital, Dallas, Texas.

Correspondence: Dr. Nasreen F. Haideri, Research Department, Texas Scottish Rite Hospital, 2222 Welborn Street, Dallas, TX 75219–3924:

Participants in the State-of-the-Science Conference on Prosthetic Foot/Ankle Mechanisms examined the body of scientific evidence that supports the clinical prescription and use of prosthetic foot and ankle mechanisms. This was accomplished through extensive review of literature comprised of clinical reports, perceptive analyses, and biomechanical studies. One consideration in reviewing literature on this subject was the lack of standardization regarding terminology and nomenclature.

Standards for nomenclature and evaluation of prosthetic devices are under development. The International Organization for Standardization (ISO) has a technical committee (TC 168) whose missions are:

  1. Establishment of a system of nomenclature and related terminology to allow all parties involved in the prosthetic/orthotic treatment of persons with disabilities to apply a standard terminology for the description of
    1. the users of prosthetic/orthotic devices
    2. the functional requirements of the devices
    3. the function of the components and the assembled devices
    4. the outcome of the delivery of the devices
  2. Establishment of a system of test methods for the verification of essential requirements on prosthetic/orthotic devices related to the safety1 of the users.

The publications that include the standards published by this technical group (Table 1) cover most terminology required to communicate clinical information regarding treatment of patients via prosthetic services (mission 1a–c, above). In addition, a substantial sum of information exists to guide the mechanical testing of prosthetic devices (mission 2). A common nomenclature to guide the design and evaluation of human subjects testing (mission 1d) is lacking. Development of this nomenclature may be slowed because of the wide variety of professionals required to reach a consensus.

Table 1

Table 1

To accurately communicate information regarding the prescription, fabrication, and evaluation of prosthetic devices, one must appreciate the disparities that can arise because of differences in adopted terminology. The current report highlights potential areas for confusion that warrant consideration, specifically clinical terminology, gait analysis, and the terminology of clinical biomechanics.

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To understand the characteristics of a prosthetic foot/ankle mechanism, it is important to relate them to a normal foot. Major joints of the human foot include the ankle joint composed of the dome of the talus and distal aspect of the tibia and fibula, the subtalar joint, the midtarsal joint and the metatarso-phalangeal joints. Accordingly, movement of the foot generally is described by motion of the ankle, hindfoot, midfoot, and forefoot. In addition to motion through major joints, the foot has a strong aponeurosis that connects the calcaneus to the forefoot and upholds an arch.

Sagittal plane plantar/dorsi flexion occurs through the ankle joint. The hindfoot generally is considered responsible for transverse and coronal plane rotations. According to Tachdjian’s textbook of orthopaedics,2 the hindfoot inverts and everts into positions of varus and valgus in the coronal plane. When the hindfoot inverts, the rest of the foot rolls onto the outer border of the foot as it supinates. When the hindfoot everts into valgus, the forefoot pronates, increasing weight bearing on the first ray. Thus, pronation and supination are movements that occur in multiple planes. The forefoot contributes to these motions by rotating along the longitudinal axis of the foot (the frontal plane). Forefoot motion in the transverse plane has been described as inversion/internal rotation or eversion/external rotation. Internal/external rotation is preferred to avoid contradiction of coronal plane movement as described for the hindfoot; however, inversion/eversion often are used in this context.

Prosthetic foot designs do not replicate exact characteristics of a normal human foot. Current prosthetic feet demonstrate some of the desired characteristics effectively but remain lacking in others.

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The primary purpose of lower extremity prosthetic devices is to facilitate ambulation. Gait is one of the most studied of all human movements. Although standardized terminology to describe events of the gait cycle do not exist, a widely used nomenclature for describing gait was developed by Perry3 (Figure 1). According to Perry, each gait cycle is divided into two periods: stance and swing. Stance is the period when the foot is in contact with the ground, constituting 62% of the gait cycle. Swing denotes the period when the foot is in the air, constituting the remaining 38% of the gait cycle. The stance period is further divided into two periods of double support occurring at the beginning and end of stance and referred to as initial and terminal double-limb stance, and a period of single-limb support when only one foot is in contact with the ground.

Figure 1.

Figure 1.

The gait cycle can be broken down in greater detail into eight phases. Initial contact represents the beginning of the stance. Loading response occupies the first 10% of the gait cycle, the period of initial double-limb support. During loading response, the foot comes in full contact with the floor, and body weight is fully transferred onto the stance limb. Midstance represents the first half of single support, 10% to 30% of the gait cycle. It begins when the contralateral foot leaves the ground (opposite toe-off) and continues as the body weight travels along the length of the foot until it is aligned over the forefoot. Terminal stance constitutes the second half of single-limb support, 30% to 50% of the gait cycle, beginning with heel rise and ending when the contralateral foot contacts the ground. During this phase, body weight moves ahead of the forefoot. Preswing is the terminal double-limb support period occupying the last 12% of stance phase (50% to 62% of the gait cycle). It begins when the contralateral foot contacts the ground and ends with ipsilateral toe-off. During this phase, the stance limb is unloaded and body weight is transferred onto the contralateral limb. Initial swing is one-third of the swing period, from 62% to 75% of the gait cycle. It begins the moment the foot leaves the ground (toe-off or terminal contact) and continues until maximum knee flexion occurs, when the swinging extremity is directly under the body and directly opposite the stance limb. Midswing occurs in the second third of the swing period, from 75% to 85% of the gait cycle. This phase begins after maximum knee flexion and ends when the tibia is in a vertical position. Terminal swing constitutes the final phase, from 85% to 100% of the gait cycle, when the tibia passes beyond perpendicular and the knee fully extends in preparation for foot contact.

The functional tasks of gait include weight acceptance, single-limb support, and limb advancement.3,4 Two phases of the stance period, initial contact and loading response, are involved in the execution of weight acceptance. This task involves initial limb stability and shock absorption while simultaneously preserving the momentum of progression as body weight is transferred onto the limb. Two phases are associated with single-limb support: midstance and terminal stance. During this task, the contralateral foot is in the swing period, and total body weight is exclusively supported on the stance limb. Stability of the stance limb as the body progresses forward is required. Four phases contribute to limb advancement: preswing, initial swing, midswing, and terminal swing. During these phases, the stance limb leaves the ground and advances forward in preparation for the next initial contact.

Gait can be analyzed by evaluating the temporal relationships of phases of the gait cycle, duration of single- and double-support periods, for example. Spatial measures of the gait cycle are also evaluated. A stride length is defined as the distance traveled from initial contact of a limb to the next initial contact of the same limb. A stride is comprised of a right and left step. Step length is measured from initial contact to contralateral initial contact (for example, the distance from right-heel contact to left-heel contact). Cadence is defined as the number of steps taken per unit of time and is the rate at which a person walks; it often is expressed in steps per minute. Velocity combines stride length and cadence and is the resultant rate of forward progression, expressed in meters per second.

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Various technologies exist to facilitate gait analysis through measurement of kinematics, movement of the body; kinetics, the forces that cause motion; electromyography, muscle activation; and metabolic energy expenditure.

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Gait kinematics, the linear and angular displacements, velocities, and accelerations, are measured by recording motion of local coordinate systems embedded in segments of the lower extremity. Measurement of kinematics relies on the ability to attach sensors to the segments of interest to establish the local embedded coordinate systems. A segment is assumed to be a rigid body, even though often multiple joints are grouped into one segment, such as in the foot. Joint motion is typically modeled with limited degrees of freedom. The degrees of freedom of motion that can be studied are related to the number of sensors used to define a segment. Both two- and three-dimensional evaluations have been reported for subjects ambulating with a prosthesis. Mathematical models are used to compute three-dimensional rotations of the pelvis, hip, knee, ankle, and foot in the sagittal, coronal, and transverse anatomical planes. The International Society for Biomechanics has proposed standards for the reporting of these joint motions5 that include motion descriptors and recommendations for segment coordinate systems to define anatomical motion using methodology extracted from Grood and Suntay.6

It is common to use models developed for normal anatomy to study prosthetic gait. This is problematic in that not all prostheses are designed with anatomically comparable joint motion. Movement occurs between adjacent parts of the device, by deformation of parts of the device, or by a combination of these. For example, prosthetic feet without ankle joint motion in the sagittal plane rely on compression of the heel cushion to facilitate foot flat. Motion of sensors located on the foot move relative to sensors located on the tibia, and this relative motion is reported as ankle plantar flexion. For this reason, kinematic evaluations of the prosthetic foot must be interpreted with some caution.

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Unless they are clearly defined, terms such as “flexion moment” or “extension moment” can be understood to mean exactly the opposite of what the speaker or writer intended.

Kinetic studies can be conducted when external forces applied to the body are measured and limb segment properties such as mass, mass moments of inertia, and center of mass locations are estimated. External forces, the ground reaction forces, typically are measured using force plates. Net joint moments are calculated via an inverse dynamics solution of the link segment model.7 Biomechanists typically report internal net joint moments, the net result of all of the internal forces acting about the joint, including moments attributable to muscles, ligaments, joint friction, and structural constraint. Expressed as an internal net joint moment, an ankle plantar flexion moment means the ankle plantarflexors (gastrocnemius, soleus) are dominant at the ankle joint, creating the moment across the joint.

Prosthetists often prefer to visualize the ground reaction force vector projected up through the body. The spatial relationship between a joint center and the ground reaction force vector influences the direction in which the joint will tend to rotate. This is a very effective tool for understanding how prosthetic componentry or alignment will affect joint stability during gait. Net joint moments viewed this way are considered external net joint moments. Consequently some confusion arises.

Consider, for example, during midstance, the force vector passes anterior to the ankle and posterior to the knee and hip. Expressing external net joint moments, the prosthetist would correctly state that a dorsiflexion moment is occurring at the ankle, a flexion moment at the knee, and an extension moment at the hip. However, the biomechanist would express internal net joint moments and correctly state that a plantar flexion moment is occurring at the ankle, an extension moment at the knee, and a flexion moment at the hip. Internal net joint moments are those required by the body to resist the external forces applied. In this example, when the ground reaction force vector passes posterior to the knee, tending to cause the knee to flex, the quadriceps or knee extensors must fire to create a knee extension moment to resist. Using internal net joint moments, the moment reflects the muscle activity that is occurring.8

Several investigators have computed net joint moments using the ground reaction force vector multiplied by the perpendicular distance from the force line of action to the joint center. This technique overestimates the net joint moments, with the errors increasing as computations are carried up the kinetic chain from the floor.4

Some investigators have chosen to report information from ground reaction force measures directly because these measures require no assumptions about the anatomy of the individual being tested or the type of joint motion occurring. Others have investigated the center of pressure trajectory plotted in a coordinate system attached to the prosthetic device, the rollover shape, which also eliminates the need for these assumptions.

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Combining kinematic and kinetic information, power can be computed as the product of net joint moment and angular velocity: P = M × ω. If P is positive, M and ω are applied in the same direction, and power is generated. Power generation is associated with concentric muscular activity. For example, a net plantar flexion moment at the ankle joint during late stance causes plantar flexion of the ankle. Therefore, the plantar flexor muscle group is exerting a concentric contraction. Likewise, during the second rocker of gait when the ankle progresses into dorsiflexion, a net plantar flexion moment is being generated, power is negative, and eccentric muscle activity occurs. Of note is that the internal net joint moment convention must be applied in these computations.

Prosthetic feet capable of storing and returning energy during gait (ESAR) have been developed. It is difficult to discern the energy storage and return capability of these devices. Testing methodologies and nomenclature to categorize prosthetic foot and ankle components based on the results of this testing are needed.

Hafner et al.9 describe one technique to evaluate performance of these devices by integration of different regions of the ankle power-time curve. Figure 2 depicts ankle power generation and absorption throughout a gait cycle. Region I constitutes an area of power absorption during loading response and region II the power generation after that phase. The ratio of these two areas describes the efficiency of the heel component of a prosthetic foot. Comparably, region III is power absorption occurring in the keel during midstance, and region IV depicts the power returned by this component at terminal stance. The ratio of these two areas can be used to describe the efficiency of the keel component of the prosthetic foot.

Figure 2.

Figure 2.

Based on this analysis, Hafner et al.9 proposed nomenclature combining four attributes to describe the function and performance of lower limb energy storage and return prosthetic devices using a single metric defining heel (H) and keel (K) energy storage and efficiency, where

  1. = high storage, high efficiency (return)
  2. = high storage, low efficiency (return)
  3. = low storage, high efficiency (return)
  4. = low storage, low efficiency (return).

Using this nomenclature, a device with good ESAR might be rated H1K1; whereas a nondynamic, low-functioning device could receive a rating of H2K4. This solution is pragmatic in that it can be applied to perceptive studies geared at extracting details regarding the subjective functioning of the foot, as well as to more quantified measures.

Since the publication of their original work, the investigators have suggested that region II, the area of energy return of the heel component, is not clinically important and often is not apparent on the power-time curve. A metric derived from a more global expression of efficiency, such as the ratio of power absorption during loading response and return during push-off, may be more relevant for clinical testing paradigms. The original more detailed nomenclature might be more appropriately applied to mechanical testing paradigms.

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Mechanical energy can be estimated using kinematic and kinetic measures, as described above. Another type of energy measured during gait analysis, metabolic energy, is a measure of the physiologic stress exerted by a body to perform the task of walking. Metabolic energy measures of gait involve sampling the amount of oxygen consumed by the body while walking to determine the intensity of exercise. These measures are typically normalized to walking velocity to determine the “cost” of ambulation per meter walked.

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Standardization of terminology and nomenclature is paramount to the communication and continued growth of the prosthetics field. Development of such standards is a veritable task because of the multidisciplinary professionals involved in the prescription, fabrication, and evaluation of prostheses. In addition, technology is advancing faster than consensus can be developed. Until standards are agreed upon, whether in the clinical setting discussing an individual patient, or in the publication and presentation of research, it is necessary to ensure that a mutual understanding of the terminology and nomenclature used is established.

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1.International Organization for Standardization, Technical Committee 168: Prosthetics and Orthotics. ISO/TC 168 Business Plan. May 2004.
2.Herring JA. Tachdjian’s Pediatric Orthopaedics From the Texas Scottish Rite Hospital for Children (3rd ed). Philadelphia: W.B. Saunders; 2002.
3.Perry J. Gait Analysis: Normal and Pathological Function. Thorofare, NJ: Slack; 1992.
4.Winter DA. Biomechanics and Motor Control of Human Gait (2nd ed). New York: Wiley; 1990.
5.Wu G, Siegler S, Allard P, et al. ISB recommendation on definitions of joint coordinate system of various joints for the reporting of human joint motion – part I: ankle, hip, and spine. J Biomech 2002;35(4):543–548.
6.Grood ES, Suntay WJ. A joint coordinate system for the clinical description of three-dimensional motions: application to the knee. J Biomech Eng 1983;105:136–144.
7.Bresler B, Frankel JP. The forces and moments in the leg during level walking. Trans Am Soc Mech Eng 1950;72:27–36.
8.Sienko Thomas S, Supan TJ. A comparison of current biomechanical terms. J Prosthet Orthot 1990;2(2):107–114.
9.Hafner BJ, Sanders JE, Czerniecki JM, Fergason J. Transtibial energy-storage-and-return prosthetic devices: a review of energy concepts and proposed nomenclature. J Rehabil Res Dev 2002;29(1):1–11.
© 2005 American Academy of Orthotists & Prosthetists