Logo for the Journal of Rehab R and D
Journal of Rehabilitation Research and Development
Vol. 38 No. 2, March/April 2001

State-of-the-art research in lower-limb prosthetic biomechanics-socket interface

Arthur F.T. Mak, PhD; Ming Zhang, PhD; David A. Boone, CP, MPH

Jockey Club Rehabilitation Engineering Centre, The Hong Kong Polytechnic University, Hong Kong, China

Address all correspondence and requests for reprints to: Professor Arthur F.T. Mak, PhD, Jockey Club Rehabilitation Engineering Centre, The Hong Kong Polytechnic University, Hong Kong, China; email:

Abstract — Scientific studies have been conducted to quantify attributes that may be important in the creation of more functional and comfortable lower-limb prostheses. The prosthesis socket, a human-machine interface, has to be designed properly to achieve satisfactory load transmission, stability, and efficient control for mobility. The biomechanical understanding of the interaction between prosthetic socket and the residual limb is fundamental to such goals. The purpose of this paper is to review the recent research literature on socket biomechanics, including socket pressure measurement, friction-related phenomena and associated properties, computational modeling, and limb tissue responses to external mechanical loads and other physical conditions at the interface. There is no doubt that improved biomechanical understanding has advanced the science of socket fitting. However, the most recent advances in the understanding of stresses experienced at the residual limb have not yet led to enough clinical consensus that could fundamentally alter clinical practice. Efforts should be made to systematically identify the major discrepancies. Further research should be directed to address the critical controversies and the associated technical challenges. Developments should be guided to offer clinicians the quantification and visualization of the interaction between the residual limb and the prosthetic socket. An understanding of comfort and optimal load transfer as patterns of socket interface stress could culminate in socket design expert systems.

Key words: amputees, biomechanics, computational biomechanics, interface, interface pressure measurement, interface shear measurement, mechanics, prosthetics, socket design, tissue mechanics, tissue responses and adaptation.


  Surveys have shown that amputees complain about their prosthesis being uncomfortable (1,2). It is not uncommon for amputees to develop skin problems on the residual limb, such as blisters, cysts, edema, skin irritation, and dermatitis (3-5). Discomfort and skin problems are usually attributed to a poor socket fit. Further improvement of prosthetic fitting is required to maximize amputee's comfort and acceptance of the prosthesis.

  The socket, as a human-device interface, should be designed properly to achieve satisfactory load transmission, stability, and efficient control for mobility. Some early designs of the prosthetic socket, such as the "plugfit," took the form of a simple cone shape, with very little rationale for the design. Previous development shows that biomechanical understanding of the interaction between the prosthetic socket and the residual limb is fundamental to the improvement of socket design. With an understanding of the residual limb anatomy and the biomechanical principles involved, more reasonable socket designs, such as the patellar tendon bearing (PTB) transtibial socket, and the quadrilateral transfemoral suction socket were developed following World War II (6,7). These designs intended to provide a more effective distribution of loads around the residual limb. These sockets are so designed that the load-tolerant areas can chiefly take the load, while relief can be given to the sensitive areas. By the 1980s, the so-called hydrostatic weight-bearing principle and the total surface bearing (TSB) concept were introduced. Examples include the silicone suction socket (8) and ICEROSS (9), as well as those incorporating the use of interfacing gel-like materials.

  The basic principles for socket design vary from either distributing most of the load over specific load-bearing areas or more uniformly distributing the load over the entire limb. No matter what kind of design, designers are interested in understanding the load-transfer pattern. This will help designers to evaluate the quality of fitting and to enhance their understanding of the underlying biomechanical rationale. Many studies have been conducted to evaluate and quantify the load distribution on the residuum by either clinical measurements or computational modeling.

  The skin and the underlying soft tissues of the residual limb are not particularly adapted to the high pressures, shear stress, abrasive relative motions, and the other physical irritations encountered at the prosthetic socket interface. In order to design a good socket fit with optimal mechanical load distributions, it is critical to understand how the residual limb tissues respond to the external loads and other physical phenomena at the interface.

  The purpose of this paper is to review the recent studies on prosthetic biomechanics, especially on the socket/residual limb interface, including 1) recent developments in socket pressure measurements, 2) recent investigations on friction-related phenomena and associated properties, such as shear stress, frictional properties of skin, slippage, et cetera, 3) computational modeling for residuum tissue stress/strain analysis, and 4) tissue responses to external mechanical loads and other physical conditions at the prosthetic interface.


  The pressure distribution at the interface between the residual limb and the prosthetic socket is a critical consideration in socket design and fit. Pressure measurements within prosthetic sockets have been conducted for about 50 years. The information obtained has been used either to increase the understanding of socket load transfer, to assess the socket design, or to validate the computational modeling.

  Interfacial pressure measurements require a proper measurement technique, including the use of transducers, their placement at the prosthetic interface, as well as the associated data acquisition and conditioning approach. An ideal system should be able to continually monitor real interfacial stresses, both pressure and shear, without significant interference to the original interface conditions. A variety of transducers have been developed for socket pressure measurements. They can be classified, based on their operation principle, as fluid-filled sensors (10-12), pneumatic sensors (13-15), diaphragm deflection strain gauge (16-25), cantilever/beam strain gauge (26-28), and printed circuit sheet sensors (29-34), as reviewed by Sanders (35) and Silver-Thorn and colleagues (36).

  The techniques for placement of transducers at the residual limb and socket interface can be divided into two categories. They are either inserted between the skin and the liner/socket, or positioned within or through the socket and/or the liner. Only thin sensors, such as the diaphragm deflection strain-gauge sensors (17,18,20,21), the fluid-filled transducers (10), the pneumatic transducers (13,15), and the printed circuit sheet sensors (29-34), are suitable for insertion between the skin and socket. Mounting is relatively easy and it is not necessary to damage the prosthesis. However, for many of these sensors, interference is unavoidable from their protrusions into the socket volume, because of their finite thickness (26,27). The diameter of each sensing element is another important consideration. Too big a sensing element can measure only an average pressure over the area, while too small a sensor may be affected by its edge effects, especially for a stiff sensor. Positioning the transducers within or through the socket with the sensing surface being flush with the skin would make the thickness of the transducer becomes less critical. For such mounting, holes would need to be made on the experimental sockets to recess the transducers (25,26,37-39).

  The techniques mentioned above can measure pressures at discrete focal sites because of the size of the sensing cells. Sensor mats with an array of pressure cells make it possible to measure the pressure distribution. However, a piece of material inserted at the interface may change the original conditions. Systems have been commercially designed for in situ socket-pressure measurements, such as the Rincoe Socket Fitting System, Tekscan F-Socket Pressure Measurement System, and Novel Pliance 16P System. The F-socket (type 9810 or 9811) transducer is a force-sensing resistor using a mylar substrate for its 0.28-mm-thick strip (40,41). There are 96 individual cells, displayed in an array of 16 rows and 6 columns, covering an area of 155 cm2. The advantages of this system are its thin and flexible sheet, acceptable sensitivity, resolution, and frequency response (42). The disadvantages usually associated with these sensors are their hysteresis, signal drift, temperature sensitivity, and unknown shear coupling effects (40-42). This system has been used for measuring the pressure distributions at socket interfaces (29,31-34). Houston and colleagues (30) reported a specially designed Tekscan P-Scan transducer with 1,360 cells. Rincoe force sensors are embedded in a polyvenilidyne fluoride strip with a thickness of 0.36 mm (41). This system has a total of 60 cells arranged on 6 separate strips, each comprised of 10 sensors. A report on the use of this system can be found in Shem and colleagues (43). The sensor pad of the Novel Pliance 16P System has 4×4 matrix capacitance sensors with 1-mm thickness. The system allows up to 16 sensor pads to be used simultaneously. There are advantages and disadvantages with each system. The performances (accuracy, hysteresis, signal drift, and the response to curvature) of the above three systems have been compared (40,41).

  The pressures reported at the socket interfaces vary widely among sites, individuals, and clinical conditions. For the PTB socket, the maximum peak pressure reportedly could reach about 400 kPa (44), the highest among all the measurements reported. However, the measurements conducted in the last 10 years showed that the maximum interface pressure for PTB sockets during walking was usually below 220 kPa (29,37,38). A wide pressure variation may result from 1) the diversity of the prostheses and fitting techniques used, 2) the difference in residual limb size, soft tissues thickness, and gait style, 3) the different positions studied, and 4) the different characteristics and limitations associated with each specific measurement and mounting method.


  The biomechanics of the coupling between the skeleton and the socket is an important factor for socket fit. This coupling is affected by the relative slippage between the subject's skin and the prosthetic socket, and the deformation of the residual limb tissues. The tightness of fit could influence the coupling stiffness. Socket shape can change the pressure distribution and the apparent tightness of fit. Generally speaking, a loose fit may allow slippage, which may compromise stability, while a very tight fit may offer a more stable connection, but increase the interface pressures. Another important factor affecting slippage is the friction between the subject's skin and the prosthetic surface. Excessive slippage at the socket interface should be avoided in socket fitting; however, absence of slippage may cause other problems. Amputees might not feel comfortable when a buffer is inserted between the skin and socket to reduce slippage (49). The discomfort apparently resulted not from the pressures, but from the increase in interface temperature and perspiration inside the socket.

  Friction is a phenomenon in which tangential force acting between bodies in contact opposes their relative motion or impending motion. Because of the existence of friction, shear forces can be applied to the skin surface. Research related to friction in the prosthetic socket includes 1) investigation of the coefficient of friction of skin with various interface materials (50-52), 2) measurements of shear stresses (37,38,45-48) and slip at the interface (53,54), 3) measurements of the relative motion between the skeleton and the prosthetic socket (55-61), and 4) the contribution of frictional shear to the load transfer.

  Frictional properties of human skin have been investigated under various skin conditions (51,62-65), to examine the effects of skin care products (66-68), and to see how friction might affect some friction-dependent manual activities (69,70). Recent studies on skin friction with various interface materials were reported by Sanders and colleagues (51) and Zhang and Mak (52). Sanders and colleagues (51) measured the coefficient of friction of in vivo human skin with eight interface materials, using a biaxial force-controlled load applicator. The measurements were conducted on shaved and cleaned skin of the lower limb. The coefficients of friction ranged from 0.48 to 0.89. The coefficients of friction with skin of the eight interface materials are significantly larger than those with sock. Zhang and Mak (52) measured the coefficient of friction of in vivo human skin with five materials, namely aluminum, nylon, silicone, cotton sock, and Pelite. The measurements were conducted on untreated skin over six anatomical sites using a Measurement Technology Skin Friction Meter. The average coefficient of friction was 0.46. The value was highest for silicone (0.61) and lowest for nylon (0.37) among the five materials studied.

  Measurements of shear stresses at the residual limb/skin interfaces were first reported by Appoldt and colleagues (45). They developed a beam deflection strain-gauge transducer, 11 mm in diameter and 27 mm in length, which could measure the normal force and shear force in one direction. Sanders and coworkers (35,37,46-48) have published a series of papers on the development of their triaxial transducers and their interface stress measurements on transtibial sockets. Shear forces in two directions were measured by mounting metal-foil strain gauges on aluminum beam. The size of the sensing surface was 6.35 mm in diameter but the gross size and weight were quite considerable (37). The transducers have been used to measure the interface stresses on the transtibial sockets to assess the shear stress magnitude (46), the transient shape of the stress waveform during walking (37), and the effects of alignment on these interface stresses (47,48). Williams and colleagues (25) developed a small size (15.9 mm in diameter and 4.9 mm in thickness) tri-axial transducer that can measure normal force and shear force in two orthogonal directions. The normal force was sensed by the diaphragm deflection strain gauges. Biaxial shear forces were sensed by magnetoresistors fixed at the center of the disk, which could slide on a cruciform to resolve the shear force into two orthogonal directions. The transducers were further used by Zhang and colleagues (38) to measure the stresses applied on the skin surface at eight locations of five transtibial sockets. A piece of Pelite material was glued on the top of the transducers fixed on the socket wall. A maximum shear stress of 61 kPa was found at the medial tibia area with PTB sockets during walking.

  Appoldt and colleagues (53) and Commean and colleagues (54) reported on the measurements of slippage between skin and prosthetic sockets. Appoldt and colleagues (53) developed a slip gage consisting of a pen rigidly held to the transfemoral sockets, whose inking tip was in light contact with the skin. The mark left on the skin was used to assess the slip magnitude and direction. The results indicated that in a well-fitted total-contact suction socket the relative slip was less than 6 mm. Commean and colleagues (54) reported the measurement of slippage between a transtibial residuum and its prosthetic socket using spiral x-ray computed tomography imaging. Lead markers were placed on the socket inner wall and skin surface. The scans were taken under two axial static loading conditions. The results showed that the relative slip increased from 2 mm to 6 mm when the applied load increased from 44.5 N to 178 N.

  The skeletal movements relative to the socket are determined by the relative slip between the skin and the prosthetic socket, as well as by the deformation of the residuum soft tissue. Radiography (55-59) and ultrasound (60,61) techniques have been used to investigate the skeletal movements within transtibial sockets (55-57) and transfemoral sockets (58-61). Using the radiographic techniques, measurements were taken under several static load conditions. The results showed that the average movement of the tibia in the proximodistal direction was 22 mm for PTB sockets (57,59) and 11 mm for PTB suction sockets (58). The use of radiographic techniques is limited by the risk of ionizing radiation and often by the static situation. Recently, ultrasound techniques were used to assess femur movements within transfemoral sockets (60,61). This technique allowed recording of such movements during walking using two simultaneously transmitting ultrasound transducers mounted on the socket. The difficulties in performing these measurements routinely in ordinary clinics are that duplicated sockets and prostheses are required, and the need for an expertise-intensive process of ultrasound data analysis (61).

  Friction between the residual limb and the prosthetic socket leads to two primary effects. Friction produces shear action on the skin and leads to tissue distortion. Such action may disturb tissue functions and can be harmful to the tissues. On the other hand, the friction-producing shear forces at the skin surface can assist in supporting the ambulatory load and in the suspension of the prosthesis during swing phase. Zhang and colleagues (33) developed an idealized cone-shaped model and a finite element (FE) model using the real limb geometry to predict the effects of friction on the load transfer. Their results showed that the smaller the friction, the smaller the shear stresses, but the larger the normal stresses required to support the same load. Experimental measurements using a Tekscan system apparently confirmed that the pressures measured at a lubricated skin/socket interface were higher than those measured at a normal residual limb/liner interface (33). Hence, reduction of interface friction may not always be a good way to alleviate residual limb tissue problems. An adequate coefficient of friction could be desirable to support loads and prevent undesirable slippage. However, a surface with large friction could experience high local stresses and tissue distortion when donning the limb into the socket, as well as during ambulation. A proper choice of friction would be needed to balance the requirements for effective prosthetic control and minimization of interfacial hazards (33).


  Although stresses at the residual limb socket interface can be measured, a full-field experimental evaluation of the load transfer remains difficult. It is anticipated that those difficulties associated with experimental measurements can be overcome by computational modeling, provided an appropriate model can be developed. With the emergence of computer-aided design and computer-aided manufacturing (CAD/CAM) technology, computational modeling is a desirable tool to provide quantitative information on the load transfer between the socket and the residual limb for the purpose of optimal socket design and objective evaluation of the fit. Computational models for socket analysis are mainly based on finite element methods. There are two major advantages in using FE analysis. First, full field information on the stress, strain, and motion anywhere within the modeled objects can be predicted. Second, it is relatively convenient to do parametric analysis for an optimal design.

  Since the computational methods were introduced to the prosthetic socket design field in the late 1980s (71,72), several FE models (73-89) have been developed, as reviewed by Zhang and colleagues (90), Silver-Thorn and colleagues (36), and Zachariah and Sanders (91). These models can be grouped into three types (90). The first type involves linear static analysis established under assumptions of linear material properties, infinitesimal deformation and linear boundary condition without considering any interface friction and slip. Models of this type require relatively small CPU time. The second type can be referred to as nonlinear analysis, taking into consideration the nonlinear material properties, large deformation, and nonlinear boundary conditions, including friction/slip contact boundary. Such nonlinear FE analysis normally requires some iterative procedures. While requiring relatively more CPU time, such nonlinear approaches generally yield more accurate solutions. The third type involves dynamic models. Analyses of this type consider not only dynamic loads, but also material inertial effects and time-dependent material properties.

  In reviewing the previous FE models, two challenges required to be addressed are 1) modeling of the residual limb soft tissues and 2) the effects of donning procedures with friction/slip interfacial conditions. Biological soft tissues, including residual limb tissues, exhibit complex mechanical properties and may undergo large deformation. The lack of an accurate description of their mechanical properties has limited the development of a precise computational model. The existing data on soft tissue properties were mainly collected using in vivo indentation tests (92-99). The material constants under the assumption of linear elasticity, isotropy, and material homogeneity were extracted by curve fitting the indentation force-deformation data with the use of FE technique (75) or by some mathematical formula transfer. The most often used mathematical model is the Hayes' solution (100), based on an elastic analysis of the infinitesimal indentation by a frictionless rigid indentor on an elastic layer bonded to a rigid foundation. The influence of friction between the indentor and the layer surface and the consideration of large deformation were included in a recent study (101). Nonlinear elastic properties, modeled as a Mooney-Rivlin material, have been used for residual tissues in some models (24).

  Simulation of donning procedure with friction/slip interfacial conditions is another challenge. In the real situation, the amputee normally puts on the liner first if fitted, and then dons the residual limb into a prosthetic socket. There are difficulties in the simulation of large displacements associated with the donning procedure. To date, most socket rectification is normally simulated by prescribing the displacement boundary conditions at the nodes on the outer surface of the socket or liner (71,75,79,80,82,84,89). Displacement boundary conditions corresponding to the shape of a given socket design are applied to deform the residual limb soft tissue or the liner to conform to the rectified socket shape. There are obvious discrepancies between such simplified simulation and the real donning procedure. Zhang and colleagues (78,79,89) applied interface elements to simulate the friction/slip boundary conditions between skin and liner. Such special four-node elements connecting the skin and the liner by corresponding nodes can be used to simulate friction and slip condition. However, they cannot be used to simulate the donning procedure when there is a large relative sliding between the liner and socket. Zachariah and Sanders (77) attempted to use an automated contact method, in which correspondence between socket and limb was not required, to simulate the friction/slip interface. Finney (102) attempted to simulate the donning by sliding the deformable residual limb into a rigid socket shell, using a simple idealized geometry.

  Further computational modeling for a residual limb and socket system can go in two primary directions. First, computational models with reliable data inputs should be further developed to become more precise, in order to better approximate the real situation. Second, computational modeling can be integrated as part of a clinical system for computer-aided socket design and manufacturing, in order to provide prosthetists with quantitative feedback during virtual socket rectification. Such clinical information must be displayed in a clinically meaningful format and the whole process would need to be user friendly.

  A complete prosthesis model should involve not only the residual limb and the socket interface, but also the whole prosthesis. Such a model can be used to discuss the effects of the prosthetic alignment, the foot/ankle joint properties, and the mass distribution of components on the load distribution between the residual limb and the socket, during the loaded support and the socket suspension phases.


  Soft tissues of the residual limb within a prosthetic socket are subjected to a special environment. First, pressures and shear forces are applied by the socket snugly fitted on the residuum, although the limb tissues are not necessarily suited for undertaking such loads. These loads are dynamic and repetitive during locomotion. Second, skin rubbing against the socket edge and interior surface may happen, resulting in intermittent skin deformation and biomechanical irritations. If excessive slip exists between the skin and the socket, tissue abrasion can occur and heat will be generated. Third, residual limb tissues exist often in a high-humidity environment, because the socket intimately fitted on residuum excludes circulating air and traps accumulated sweat. Fourth, the residual limb tissues may suffer from possible chemical and mechanical irritations or allergic reactions to various socket or interface materials (4,103).

  Under such an unfavorable and demanding environment, whether the residual limb soft tissues will break down or adapt is a primary consideration in socket design and fitting. If a good skin condition cannot be maintained, the prosthesis can no longer be worn, no matter how accurate the fit of the socket may be. Clinical treatment and socket fit should encourage skin adaptation and avoid breakdown (104). The following review focuses on the response of soft tissues to external loads. The literature reviewed includes studies not only on residual limb tissues, but also other soft tissues that are in contact with external supporting surfaces.

  Tissue responses to external forces are complicated, involving tissue deformation, interstitial fluid flow, ischemia, reactive hyperemia, sweat, pain, skin temperature, skin color, et cetera. In general, normal physiological forces will not normally disrupt tissue functions. However, an improper application either of an unusually very large force or of a prolonged or repetitive force may damage functions and/or structures. Mechanically, forces applied to skin surface will produce stresses and strain within the skin and the underlying tissues. Those stresses and deformation affect cellular functions and other biophysical processes in the tissues. A very large force may break the skin directly. When moderate static forces are applied to the skin, the underlying blood vessels and lymphatic drainage can be occluded or partially occluded, and oxygen and other nutrients can no longer be delivered at a rate sufficient to satisfy the metabolic requirements of the tissues. Without a sufficient circulation, the breakdown products of metabolism would accumulate within the tissues. If such a condition continues, cellular functions would be compromised and could ultimately fail (105). Tissue breakdown occurs not only on the skin surface, but is often found also in deep tissues (108,109).

  Repetitive forces may damage the tissues by accumulating their effects. Although a moderate force may not cause direct and immediate damage to the tissues, repeated applications day after day could initiate an inflammation reaction, and even result in tissue necrosis. When the applied load is within certain windows, tissue adaptation may occur by changing its tissue composition and architecture (106,107).

  Besides the load magnitude, other load characteristics, such as direction, distribution, duration (3Ds), and loading rate should be considered in the discussion of soft tissue responses to external loads. The forces applied to the skin surface can be resolved into two components, normal force perpendicular to the skin surface and shear force tangential to skin surface. Some researchers (110,111) suggested that tissue deformation or distortion, rather than mere pressures, are important variables in the study of tissue damage by external loads. When the pressures are evenly distributed over a wide area of the body, damage is apparently less than when loading is applied over a localized area (112). It is generally agreed that an inverse relationship exists between the intensity of the external loads and the duration of load application required to produce ulceration (108,113-115). A number of studies have been presented to theoretically explain such an inverse relationship (116-119). Mak and coworkers (118,119) invoked the physics of interstitial fluid flows induced by a given epidermal pressure to account for the corresponding endurance time. Landsman and colleagues (120) hypothesized that a higher strain rate of tissue deformation may cause a higher pressure buildup in the tissues and a higher elevation of intracellular calcium concentration, leading potentially to more damage to the involved tissues.

  Pain, or discomfort, is the most direct reaction of the human body to excessive external loads. When an abnormally large force is applied to a skin surface, the subject will normally feel some level of pain immediately. Normal sensory function of a human body can often help to avoid a mechanical insult and the subsequent tissue damage. Such sensory feedback can prompt the subject to stop or avoid further application of the loads. Neuropathy can lead to the loss of this function and may result in otherwise preventable damage, such as in the formation of pressure ulcers in diabetic and spinal cord-injured patients.

  Load-related thresholds for pain vary with anatomical locations and from person to person. Investigations have been performed to measure the ability of the human body to sustain external forces. The general measurements involve the pressure threshold, i.e., the minimum pressure to induce pain or discomfort, and the pressure tolerance, i.e., the maximum pressure a person can tolerate without excessive effort (121). For residual limbs, the tolerant and sensitive areas have been identified qualitatively (6). Studies have been reported on the load-tolerance levels of the distal ends of residual limbs (122,123).

Microvascular Responses
  It is generally believed that ischemia is related to the formation of pressure sores. Ischemia can lead to local malnutrition. Changes in local skin blood supply under various external loading conditions have been studied for a number of years. A series of reports have described the effects of external loads on skin blood flow using radionuclide clearance (124-126), photoplethysmography (127,128), transcutaneous oxygen tension (129-131), and laser Doppler flowmetry (132-139). The results of these studies seem to show that the blood supply would be influenced by the epidermal forces, and the rate and the amount of blood supply would decrease with increased epidermal loads.

  Investigations have been done to understand the effects of shear forces in conjunction with normal forces (127,136-138,140). It was noted that cutaneous blood flow was reduced with the increased application of either the normal force or the shear force. The resultant force is a critical parameter in assessing the combined effect of these multi-axial loads (137). Tam and colleagues (138) compared the reactive hyperemia in skin induced by the application of a normal force and that due to the application of both normal and shear forces. It was found that the addition of shear force would increase the tissue recovery time from the effects of hyperemia. This recovery time was taken as indicative of the tissue capacity to accommodate the biomechanical challenges.

Lymphatic Supply and Metabolites
  The lymphatic system consists of a complex network of vessels, and presents a drainage route for the transport of excess fluid, protein, and metabolic wastes from the tissue of origin into the circulatory system. External loads may interfere with the normal function of this system. With tissue edema, poor lymphatic function was associated with sore formation (112). Krouskop and colleagues (141) suggested that the smooth muscle of the lymphatics was sensitive to anoxia, and thus the impairment of the lymphatic function combined with changes in the microvascular system could compromise tissue viability through the accumulation of metabolic wastes.

  The levels of metabolites in sweat may be used as indicators of the tissue viability status (142,143). Studies showed that epidermal loads could change the amounts and the composition of sweat (144). It was found that there was a significant increase in sweat lactate during loading and a decrease in sweat volume during ischemia.

Skin Temperature
  Skin temperature may be taken as a stress indicator for the tissues (145). It was hypothesized that reduced blood perfusion during load application on the skin would be expected to lead to a local fall in skin temperature, and a rise in temperature is expected with the subsequent reactive hyperemia upon load removal. It was suggested that this temperature information might become useful during prosthetic fitting as indicative of the local pressure distribution (146,147). It was shown that tissue temperature decreased as a direct consequence of applied loads (135,145,148,149).

  However, whether the temperature can be an indicator of tissue problems is still arguable. Skin temperature is influenced by many factors, which can readily interfere with the absolute surface temperature measurement (145). Schubert and Fagrell (150) found that the temperature increased by 2.7°C over the gluteus and 1.3°C over the sacrum when a repetitive normal force was applied to those areas. The contact materials may accumulate heat, which may affect the tissue temperature response. It has been remarked that increasing the temperature by 1°C can have an effect of increasing the metabolic demands of the cells and oxygen consumption by 10 percent in the associated tissue (151). Skin blood perfusion rate is related to the environmental temperature. Blood flow would increase with a warming environment (152,153).

  An increase in temperature may be an early signal for the formation of pressure sores (154). However, apparent controversies still exist. The mean foot temperature of the painful diabetic neuropathic patients is significantly higher than that of the control subjects (155,156). The temperature of the skin at risk of pressure sores was not found to be higher than the other healthy areas (157), although the rate of blood supply markedly differed.

Skin Abrasion
  Frictional rubbing is one of the most common insults to which the human skin is exposed (158). It can produce a variety of skin lesions such as calluses, corns, thickening, abrasions, and blisters (159). Repetitive rubbing produces heat, which may cause uncomfortable and detrimental consequence (103). Naylor (158) summarized two kinds of skin reactions to repeated rubbing. One involved skin thickening if the abrasive force is small but rubbing is frequently repeated. The other involved the formation of blisters if the abrasive force is large. It was observed that blisters apparently do not often form on thin skin, but on tough and thick skin (159). Experiments have been conducted to study skin lesions under repetitive pressure with and without the involvement of frictional force (159-161). Results indicated that the addition of friction would accelerate skin damage. Sanders (162) measured the thermal response of skin to cyclic pressure alone and to cyclic pressure plus shear. The results from three normal subjects indicated that the thermal recovery time was apparently higher for combined pressure and shear compared to the values for pressure alone. The apparent additional insults due to shear as demonstrated in this study were consistent with other skin perfusion studies (138).


  The fundamental goal of prosthetic interface biomechanics research is to achieve optimal and not merely adequate function. Even the most rigorous scientific analyses to date have focused in large part on socket designs based on historical use and proven clinical adequacy. Instrumentation and computer modeling have been useful in illuminating what had only previously been the implied conditions inside prosthetic sockets. However, the most recent advances in the understanding of stresses experienced at the limb/prosthesis interface have not yet fundamentally altered clinical practice (163). Still, it is increasingly necessary for clinicians to cope with new prosthesis designs and materials that do not have the benefit of long histories of successful application. For example, use of new materials such as elastomeric liners and flexible thermoplastic sockets necessarily alter the manner in which load is transferred from the limb to the prosthesis. Improved understanding of prosthetic interface stresses allows us to understand the biomechanical effect of these new interfaces and can help prosthetists to adjust their socket designs to make best use of the properties of new technologies.

  For all prosthetic socket designs, the optimal load distribution should be proportional to the ability of the body to sustain such stresses, without crossing the thresholds of pain or skin breakdown. More research is required to obtain sufficient quantitative data to fully document these tissue threshold properties and their dependence on age and pathologies. Without a rigorous understanding of these tissue properties, it would be futile to discuss optimal load distribution and how to achieve that by various prosthetic socket designs.

  The CAD/CAM technology for the prosthetic socket may make the socket design and manufacture process more effective and objective. However, the current CAD/CAM systems cannot offer any expert suggestion on how to make an optimal socket design. Further improvement of the systems should incorporate qualification and visualization of the interaction between the residual limb and the prosthetic socket. Computational modeling with further improvements can be a useful tool for this purpose. If the research can accumulate enough information on the relationship between quantified values and the comfort of the prosthesis, CAD/CAM systems can be further developed into expert systems that propose an optimal socket configuration.

  The most radical of new prosthetic developments is certainly direct skeletal attachment of limb prostheses through osseointegrated implants. This method completely obviates the need for the prosthetic socket through percutaneous titanium fixtures that transfer load from the prosthesis directly to the skeletal bone. While it may seem that osseointegration renders moot any discussion of prosthetic interfaces, even this radical advance in the state of the art only changes the location and type of the interface problem. New challenges arise from the metal/bone and metal/skin interfaces. The latter juncture is of particular importance because it must artificially provide the critical skin barrier to the environment. Responses of soft tissues to the abnormal stresses at the point of attachment are somewhat related to interface mechanics studied previously.

  Prosthetic biomechanics is one of the most challenging areas in the field of biomechanics. There is no doubt that improved biomechanical understanding has advanced the science of socket fitting. However, the most recent advances in the understanding of stresses experienced at the residual limb have not yet led to enough clinical consensus that could fundamentally alter clinical practice. Efforts should be made to systematically identify the major discrepancies. Further research should be directed to address the critical controversies and the associated technical challenges. To these ends, we hope this review article could offer some contribution.

  1. McColl I. Review of artificial limb and appliance centre services. London: Department of Health and Social Security (DHSS); 1986.
  2. Nielsen CC. A survey of amputees: functional level and life satisfaction, information needs, and the prosthetist's role. J Prosthet Orthot 1990;3:125-9.
  3. Allende MF, Levy SW, Barnes GH. Epidermoid cysts in amputees. Acta Dermato-Venereologica 1963;43:56-67.
  4. Levy SW. Skin problems of the leg amputee. Prosthet Orthot Int 1980;4:37-44.
  5. Lyon CC, Kulkarni J, Zimerson E, Van Ross E, Beck MH. Skin disorders in amputees. J Am Acad Dermatol 2000;42:501-7.
  6. Radcliff CW, Foort J. The patellar-tendon-bearing below-knee prosthesis. Berkeley, CA: Biomechanics laboratory, University of California; 1961.
  7. Radcliff CW. Functional considerations in the fitting of above-knee prostheses. Artificial Limb 1955;2(1):35-60.
  8. Fillauer CE, Pritham CH, Fillauer KD. Evaluation and development of the silicone suction socket (3S) for below-knee prostheses. J Prosthet Orthot 1989;1:92-103.
  9. Kristinsson O. The ICEROSS concept: a discussion of a philosophy. Prosthet Orthot Int 1993;17:49-55.

    Pressure and Shear Measurements

  10. Van Pijkeren TV, Naeff M, Kwee HH. A new method for measurement of normal pressure between amputation residual limb and socket. Bull Prosthet Res 1980;17(1):31-4.
  11. Naeff M, Van Pijkeren TV. Dynamic pressure measurements at the interface between residual limb and socket--the relationship between pressure distribution, comfort, and brim shape. Bull Prosthet Res 1980;10-33:35-50.
  12. Isherwood PA. Simultaneous PTB socket pressures and force plate values. BRADU Report, London: Biomechanical Res and Dev Unit; 1978. p. 45-9.
  13. Mueller SJ, Hettinger T. Measuring the pressure distribution in the socket of the prostheses. Orthpadie-Technik Heft 1954;9:222-5.
  14. Lebiedowsky N, Kostewicz J. Determination of the order of pressure exerted by static forces on the skin of the lower limb stump with the prosthesis. Chir Narz, Ruchu Orop polska 1977;42:615-24.
  15. Kroupskop TA, Brown J, Goode B, Winningham D. Interface pressures in above-knee sockets. Arch Phys Med Rehabil 1987;68:713-4.
  16. Sonck WA, Cockrell JL, Koepke GH. Effect of linear material on interface pressures in below-knee prostheses. Arch Phys Med Rehabil 1970;51(11):666-9.
  17. Rae JW, Cockrell JL. Interface pressure and stress distribution in prosthetic fitting. Bull Prosth Res 1971;10-16:64-111.
  18. Pearson JR, Holmgren G, March L, Oberg K. Pressure in critical regions of the below-knee patellar-tendon-bearing prosthesis. Bull Prosth Res 1973;10-19:52-76.
  19. Pearson JR, Holmgren G, March L, Oberg K. Pressure variation in the below-knee patellar tendon bearing suction socket. J Biomech 1974;7:487-96.
  20. Burgess EM, Moor AJ. A study of interface pressure in the below-knee prosthesis (Physiological suspension: an interim report). Bull Prosthet Res 1977;Fall:58-70.
  21. Winaski D, Pearson JR. Least-squares matrix correlations between stump stresses and prosthesis loads for below-knee amputees. J Biomech Eng 1987;109:238-46.
  22. Leavitt LA, Petersson CR, Canzoneri J, Pza R, Muilenburg AL, Rhyne VT. Quantitative method to measure the relationship between prosthetic gait and the forces produced at the stump-socket interface. Am J Phys Med 1970;49:192-203.
  23. Leavitt LA, Zuniga EN, Calvert JC, Canzoneri J, Petersson CR. Gait analysis and the tissue-socket interface pressures in above-knee amputees. Southern Med J 1972;65:1197-207.
  24. Steege JW, Childress DS. Finite element modeling of the below-knee socket and limb: phase II. Modelling and Control Issues in Biomechanical System Symposium, ASME WAM, BED, 1988;11:121-9.
  25. Williams RB, Porter D, Roberts VC, Triaxial force transducer for investigating stresses at the stump/socket interface. Med Biol Eng Comput 1992;1:89-96.
  26. Appoldt FA, Bennett L. A preliminary report on dynamic socket pressure. Bull Prosthet Res 1967;Fall:20-55.
  27. Appoldt FA, Bennett L, Contini R. Socket pressure as a function of pressure transducer protrusion. Bull Prosthet Res 1969;Spring:236-49.
  28. Sanders JE, Daly CH. Measurement of stresses in the three orthogonal directions at the residual limb-prosthetic socket interface. IEEE Trans Rehabil Eng 1993;1(2):79-85.
  29. Engsberg JR, Springer MJN, Harder JA. Quantifying interface pressure in below-knee-amputee socket. J Assoc Children's Prosthetic-Orthotic Clinics 1992;27(3):81-8.
  30. Houston VL, Mason CP, LaBlanc KP, Beatties AC, Garbarini MA, Lorenze EJ. Preliminary results with the DVA-Tekscan BK prosthetic socket/residual limb stress measurement system. Proceedings of 20th Annual Meeting American Academy of Orthotist and Prosthetist, Nashville, TN, 1994; p. 8-9.
  31. Covery P, Buis AWP. Conventional patellar-tendon-bearing (PTB) socket/stump interface dynamic pressure distribution recorded during the prosthetic stance phase of gait of a trans-tibial amputee. Prosthet Orthot Int 1998;22:193-8.
  32. Covery P, Buis AWP. Socket/stump interface dynamic pressure distribution recorded during the prosthetic stance phase of gait of a trans-tibial amputee wearing a hydrocast socket. Prosthet Orthot Int 1999;23:107-12.
  33. Zhang M, Turner-Smith AR, Tanner A, Roberts VC. Frictional action at residual limb/prosthetic socket interface. Med Eng Phys 1996;18(3):207-14.
  34. Zhang M, Mak AFT, Chung AKI. Dynamic pressure maps over areas of AK prosthetic sockets. Proceedings of 9th Word Congress of ISPO, Amsterdam, 1998; p. 709-11.
  35. Sanders JE. Interface mechanics in external prosthetics: review of interface stress measurement techniques. Med Biol Eng Comput 1995;33:509-16.
  36. Silver-Thorn MB, Steege JW, Childress DS. A review of prosthetic interface stress investigations. J Rehabil Res Dev 1996;33(3):253-66.
  37. Sanders JE, Daly CH, Burgess EM. Clinical measurement of normal and shear stresses on a trans-tibial stump: characteristics of wave-form shapes during walking. Prosthet Orthot Int 1993;17:38-48.
  38. Zhang M, Turner-Smith AR, Tanner A, Roberts VC. Clinical investigation of the pressure and shear stress on the trans-tibial stump with a prosthesis. Med Eng Phys 1998;20(3):188-98.
  39. Lee VSP, Solomonidis SE, Spence WD. Stump-socket interface pressure as an aid to socket design in prostheses for trans-femoral amputees--a preliminary study. J Eng Med 1997;211:167-80.
  40. Polliack AA, Sieh RC, Craig DD, Landsberger S, McNeil DR, Ayyappa E. Scientific validation of two commercial pressure sensor systems for prosthetic socket fit. Prosthet Orthot Int 2000;24:63-73.
  41. Polliack AA, Landsberger S, McNeil DR, Sieh RC, Craig DD, Ayyappa E. Socket measurement systems perform under pressure. Biomechanics 1999;June:71-80.
  42. Buis AWP, Covery P. Calibration problems encountered while monitoring stump/socket interface pressures with force sensing resistors: techniques adopted to minimise inaccuracies. Prosthet Orthot Int 1997;21:179-82.
  43. Shem KL, Breahey JW, Werner PC. Pressures at the residual limb-socket interface in trans-tibial amputees with thigh lacer--side joints. J Prosthet Orthot 1998;10:51-5.
  44. Meier RH, Meeks ED, Herman RM. Stump-socket fit of below-knee prostheses: comparison of three methods of measurement. Arch Phys Med Rehabil 1973;54:553-8.
  45. Appoldt FA, Bennett L, Contini R. Tangential pressure measurement in above-knee suction sockets. Bull Prosthet Res 1970;Spring:70-86.
  46. Sanders JE, Daly CH, Burgess EM. Interface shear stresses during ambulation with a below-knee prosthetic limb. J Rehabil Res Dev 1992;29(4):1-8.
  47. Sanders JE, Daly CH. Interface pressure and shear stresses: sagittal plane angular alignment effects in three trans-tibial amputee case studies. Prosthet Orthot Int 1999;23:21-9.
  48. Sanders JE, Bell DM, Okumura RM, Dralle AJ. Effects of alignment changes on stance phase pressures and shear stresses on transtibial amputees: measurements from 13 transducer sites. IEEE Trans Rehabil Eng 1998;6(1):21-31.

    Friction, Slip, and Skeleton Movements

  49. Allende MF, Levy SW, Barnes GH. Epidermoid cysts in amputees. Acta Dermato-Venereologica 1963;43:56-67.
  50. Naylor PFD. The skin surface and friction. Br J Derm 1955;67:239-48.
  51. Sanders JE, Greve JM, Mitchell SB, Zachariah SG. Material properties of commonly-used interface materials and their static coefficients of friction with skin and socks. J Rehabil Res Dev 1998;35(2):161-76.
  52. Zhang M, Mak AFT. In vivo friction properties of human skin. Prosthet Orthot Int 1999;23:135-41.
  53. Appoldt FA, Bennett L, Contini R. The results of slip measurements in above-knee suction sockets. Bull Prosthet Res 1968;Fall:106-12.
  54. Commean PK, Smith KE, Vannier MW. Lower extremity residual limb slippage within the prosthesis. Arch Phys Med Rehabil 1997;78(5):476-85.
  55. Eriksson U, Lemperg R. Roentgenological study of movements of the amputation stump within the prosthesis socket in below-knee amputees fitted with a PTB prosthesis. Acta Orthop Scand 1969;40:520-9.
  56. Grevsten S, Erikson U. A roentgenological study of the stump-socket contact and skeletal displacement in the PTB-suction prosthesis. Upsala J Med Sci 1975;80:49-57.
  57. Lilja M, Johnsson T, Öberg T. Movement of the tibial end in a PTB prosthesis socket: a sagittal X-ray study of the PTB prosthesis. Prosthet Orthot Int 1993;17:21-6.
  58. Long I. Normal shape normal alignment (NSNA) above-knee prosthesis. Clin Prosthet Orthot 1985;29(4):53-4.
  59. Sabolich J. Contoured adduction trochanteric controlled alignment method (CAT-CAM). Clin Prosthet Orthot 1985;9(4):15-26.
  60. Convery P, Murray KD. Ultrasound study of the motion of the residual femur within a trans-femoral socket during gait. Prosthet Orthot Int 2000;24:226-32.
  61. Murray KD, Convery P. The calibration of ultrasound transducers used to monitor motion of the residual femur within a trans-femoral socket during gait. Prosthet Orthot Int 2000;24:55-62.
  62. Loden M. Biophysical properties of dry atopic and normal skin with special reference to effects of skin care products. ACTA Dermato-Venereologica 1995;192(Suppl):3-48.
  63. Elsner P, Wilhelm D, Maibach HI. Frictional properties of human forearm and vulvar skin: influence of age and correlation with transepidermal water loss and capacitance. Dermatological 1990;181:88-91.
  64. Comaish JS, Bottoms E. The skin and friction: deviations from amonoton's laws, and the effects of hydration and lubrication. Br J Derm 1971;84:37-43.
  65. Cua AB, Wilhelm KP, Maibach HI. Frictional properties of human skin: relation to age, sex and anatomical region, stratum corneum hydration and transepidermal water loss. Br J Derm 1990;123:473-9.
  66. El-Shimi AF. In vivo skin friction measurements. J Soc Cosmet Chem 1977;28:37-51.
  67. Wolfram LJ. Friction of skin. J Soc Cosmet Chem 1983;34:465-76.
  68. Nacht S, Close JA, Yeung D, Gans EH. Skin friction coefficient: changes induced by skin hydration and emollient application and correlation with perceived skin feel. J Soc Cosmet Chem 1981;32:55-65.
  69. Johansson RS, Cole KJ. Grasp stability during manipulative actions. Can J Physiol Pharmacol 1994;72:511-24.
  70. Buchholz B, Frederick LJ, Armstrong TJ. An investigation of human palmar skin friction and the effects of materials, pinch force and moisture. Ergonomics 1988;31(3):317-25.

    Computational Modeling

  71. Krouskop TA, Muilenberg AL, Doughtery DR, Winningham DJ. Computer-aided design of a prosthetic socket for an above-knee amputee. J Rehabil Res Dev 1987;24(2):31-8.
  72. Steege JW, Schnur DS, Vorhis RL, Rovick JS. Finite element analysis as a method of pressure prediction at the below-knee socket interface. Proceedings of RESNA 10th Annual Conference, California; 1987. p. 814-6.
  73. Steege JW, Childress DS. Finite element prediction of pressure at the below-knee socket interface. Report of ISPO Workshop on CAD/CAM in Prosthetics and Orthotics; 1988. p. 71-82.
  74. Childress DS, Steege JW. Computer-aided analysis of below-knee socket pressure. J Rehabil Res Dev Progress Report. 1987;2:22-4.
  75. Reynolds DP, Lord M. Interface load analysis for computer-aided design of below-knee prosthetic sockets. Med Biol Eng Comput 1992;30:419-26.
  76. Sanders JE, Daly CH. Normal and shear stresses on a residual limb in a prosthetic socket during ambulation: comparison of finite element results with experimental measurements. J Rehabil Res Dev 1993;30(2):191-204.
  77. Zachariah SG, Sanders JE. Finite element estimates of interface stress in the trans-tibial prosthesis using gap elements are different from those using automated contact. J Biomech 2000;33:895-9.
  78. Zhang M, Roberts VC. Development of a nonlinear finite element model for analysis of stump/socket interface stresses in below-knee amputee. In: Held KD, Brebbia CA, Ciskowski RD, Power H, editors. Computational biomedicine. Computational Mechanics Pub, Southampton; 1993. p. 209-14.
  79. Zhang M, Lord M, Turner-Smith AR, Roberts VC. Development of a nonlinear finite element modelling of the below-knee prosthetic socket interface. Med Eng Phys 1995;17(8):559-66.
  80. Silver-Thorn MB, Childress DC. Parametric analysis using the finite element method to investigate prosthetic interface stresses for persons with trans-tibia amputation. J Rehabil Res Dev 1996;33(3):227-38.
  81. Quesada P, Skinner HB. Analysis of a below-knee patellar tendon-bearing prosthesis: a finite element study. J Rehabil Res Dev 1991;3:1-12.
  82. Brennan JM, Childress DS. Finite element and experimental investigation of above-knee amputee limb/prosthesis systems: a comparative study. ASME Advances in Bioengineering 1991;20:547-50.
  83. Torres-Moreno R, Jones D, Solomonidis SE, Mackie H. Magnetic resonance imaging of residual soft tissues for computer-aided technology applications in prosthetics--a case study. J Prosthet Orthot 1999;11(1):6-11.
  84. Torres-Moreno R, Solomonidis SE, Jones D. Three-dimensional finite element analysis of the above-knee residual limb. Proceedings of 7th World Congress ISPO; 1992. p. 274.
  85. Seguchi Y, Tanaka M, Nakagawa A, Kitayama I. Finite element analysis and load identification of above-knee prosthesis socket. Proceedings of 4th Int ANSYS Conf Exhib, Pittsburgh; 1989. p. 12.31-12.44.
  86. Krouskop TA, Dougherty DR, Vinson FS. A pulsed Doppler ultrasonic system for making noninvasive measurements of the mechanical properties of soft tissue. J Rehabil Res Dev 1987;2:1-8.
  87. Mak AFT, Yu YM, Hong LM, Chan C. Finite element models for analyses of stresses within above-knee stumps. Proceedings of 7th ISPO, Chicago; 1992. p. 147.
  88. Lee VSP, Solomonidis SE, Spence WD, Paul JP. A study of the biomechanics of the residual limb/prosthesis interface in trans-femoral amputees. Proceedings of 8th Word Congress of ISPO, Melbourne, Australia; 1994. p. 79.
  89. Zhang M, Mak AFT. A finite element analysis of the load transfer between an above-knee residual limb and its prosthetic socket - Roles of interfacial friction and distal-end boundary conditions. IEEE Trans Rehabil Eng 1996;4(4):337-46.
  90. Zhang M, Mak AFT, Roberts VC. Finite element modelling of a residual lower-limb in a prosthetic socket--a survey of the development in the first decade. Med Eng Phys 1998;20(5):360-73.
  91. Zachariah SG, Sanders JE. Interface mechanics in lower-limb external prosthetics: a review of finite element models. IEEE Trans Rehabil Eng 1996;4(4):288-302.
  92. Mak AFT, Liu GHW, Lee SY. Biomechanical assessment of below-knee residual limb tissue. J Rehabil Res Dev 1994;31(3):188-98.
  93. Malinauskas M, Krouskop TA, Barry PA. Noninvasive measurement of the stiffness of tissue in the above-knee amputation limb. J Rehabil Res Dev 1989;3:45-52.
  94. Vannah WM, Childress DS. Indentor tests and finite element modeling of bulk muscular tissue in vivo. J Rehabil Res Dev 1996;33(3):239-52.
  95. Vannah WM, Childress DS. Modelling the mechanics of narrowly contained soft tissues: the effects of specification of Poisson's ratio. J Rehabil Res Dev 1993;30(2):205-9.
  96. Vannah WM, Drvaric DM, Hastings JA, Stand JA, Harning DM. A method of residual limb stiffness distribution measurement. J Rehabil Res Dev 1999;36(1):1-7.
  97. Zheng YP, Mak AFT. An ultrasound indentation system for biomechanical properties assessment of soft tissues in-vivo. IEEE Trans Biomed Eng 1996;43(9):912-8.
  98. Zheng Y, Mak AFT. Effective elastic properties for lower limb soft tissues from manual indentation experiment. IEEE Trans Rehabil Eng 1999;7(3):257-67.
  99. Zheng YP, Mak AFT, Lue BK. Objective assessment of limb tissue elasticity: development of a manual indentation procedure. J Rehabil Res Dev 1999;36(2):71-85.
  100. Hayes WC, Keer LM, Herrmann G, Mockros LF. A mathematical analysis for indentation tests of articular cartilage. J Biomech 1972;5:541-51.
  101. Zhang M, Zheng YP, Mak AFT. Estimating the effective Young's modulus of soft tissues from indentation tests. Med Eng Phys 1997;19(6):512-7.
  102. Finney L. Simulation of donning a prosthetic socket using an idealised finite element model of the residual limb and prosthetic socket. Proceedings of 10th Int Conf Biomed Eng, Singapore; 2000. p. 71.

    Tissue Responses to Mechanical Loading

  103. Barnes G. Skin health and stump hygiene. Artificial Limbs 1956;3(1):4-19.
  104. Sanders JE, Goldstein BS, Leotta DF. Skin response to mechanical stress: adaptation rather than breakdown--a review of the literature. J Rehabil Res Dev 1995;32(3):214-26.
  105. Michel CC, Gillott H. Microvascular mechanisms in stasis and ischaemia. In: Bader DL, editor. Pressure sores--clinical practice and scientific approach. London: MacMillan Press; 1990. p. 153-63.
  106. Mackenzie IC. The effects of frictional stimulation on mouse ear epidermis. J Inv Dermatol 1974;62:80-5.
  107. Baker SR. Fundamentals of expanded tissue. Head Neck 1991;13:327-33.
  108. Daniel RK, Priest DL, Wheatley DC. Etiologic factors in pressure sores: an experimental model. Arch Phys Med Rehabil 1981;61:492-8.
  109. Groth KE. Clinical observations and experimental studies on the origin of decubiti. Acta Chir Scand 1942;(Suppl)76(1):1-209.
  110. Gillott H. Quantifying the fit of a hand orthosis. Care Brit J Rehabil Tissue Viab 1985;1(1):12-7.
  111. Levine N. Friction blisters. Physician and Sports medicine 1982;10(3):84-92.
  112. Husain T. An experimental study of some pressure effects on tissues with reference to the bed-sore problem. J Path Bact 1953;66:347-58.
  113. Kosiak M. Etiology and pathology of ischemic ulcers. Arch Phys Med Rehabil 1959;39:62-9.
  114. Reswick JB, Rogers JE. Experience at Rancho Los Amigos Hospital with device and techniques to prevent pressure sores. In: Kenedi RM, Cowden JM, Scales JT, editors. Bed sore biomechanics. London: MacMillan Press; 1976. p. 301-10.
  115. Akbarzadeh MR. Behaviour for relieving pressure. In: Webster JG, editor. Prevention of pressure sores, engineering and clinical aspects, Bristol: IOP Pub Ltd; 1991. p. 175-90.
  116. Reddy NP, Cochran GVB, Krouskop TA. Interstitial fluid as a factor in decubitus ulcer formation. J Biomech 1981;14(2):879-81.
  117. Sacks AH. Theoretical prediction of a time-at pressure curve for avoiding pressure sores. J Rehabil Res Dev 1989;26(3):27-34.
  118. Mak AFT, Huang LD, Wang Q. A biphasic poroelastic analysis of the flow dependent subcutaneous tissue pressure and compaction due to epidermal loading: issues in pressure sores. Trans ASME J Biomech Eng 1994;116:421-9.
  119. Zhang JD, Mak AFT, Huang LD. A large deformation biomechanical model for pressure ulcers. Trans ASME J Biomech Eng 1997;119:406-8.
  120. Landsman AS, Meaney DF, Cargill RS, Macarak EJ, Thibault LE. High strain rate tissue deformation. J Am Podiatric Med Assoc 1995;85(10):519-27.
  121. Fischer AA. Pressure tolerance over muscles and bones in normal subjects. Arch Phys Med Rehabil 1986;67:406-9.
  122. Persson BM, Liedberg E. Measurement of maximum end-weight-bearing in lower limb amputees. Prosthet Orthot Int 1982;6:147-51.
  123. Katz K, Susak Z, Seliktar R, Najenson T. End-bearing characteristics of patellar-tendon-bearing prostheses--a preliminary report. Bull Prosth Res 1979;BPR10-32:55-68.
  124. Daly CH, Chimoskey JE, Holloway GA, Kennedy D. The effect of pressure loading on the blood flow rate in the human skin, In: Kenedi RM, Cowden JM, Scales JT, editors. Bed sore biomechanics. London: MacMillan Press; 1976. p. 69-77.
  125. Larsen B, Holstein P, Lassen NA. On the pathogenesis of pressure sores, skin blood flow cessation by external pressure on the back. Scand J Plas Reconstr Surg 1979;13:347-50.
  126. Romanus M, Stenqvist O, Svensjo E, Ehira T. Microvascular changes due to repeated local pressure-induced ischaemia: intravital microscope study on hamster cheek pouch. Arch Phys Med Rehabil 1983;64:553-5.
  127. Bennett L, Kavner D, Lee BK, Frieda A. Shear vs pressure as causative factors in skin blood flow occlusion. Arch Phys Med Rehabil 1979;60:309-14.
  128. Bennett L, Kavner D, Lee BY, Trainor FS, Lewis J. Skin stress and blood flow in sitting paraplegic patients. Arch Phys Med Rehabil 1984;65:186-90.
  129. Newson TP, Pearcy MJ, Rolfe P. Skin surface PO2 measurement and the effect of externally applied pressure. Arch Phys Med Rehabil 1981;62:390-2.
  130. Bader DL, Gant CA. Changes in transcutaneous oxygen tension as a result of prolonged pressers at the sacrum. Clin Phys Physiol Meas 1988;9(1):33-40.
  131. Sangeorzan BJ, Harrington RM, Wyss CR, Czerniecki JM, Matsen FA. Circulatory and mechanical response of skin to loading. J Orthop Res 1989;6(30):425-31.
  132. Ek AC, Gustavsson G, Lewis DH. Skin blood flow in relation to external pressure and temperature in the supine position on a standard hospital mattress. Scand J Rehabil Med 1987;19:121-6.
  133. Sableman EE, Valainis E, Sacks AH. Skin capillary blood flow under very light pressure. Proc Ann Conf Rehabil Eng Soc N Am 1989;12:256-7.
  134. Sacks AH, O'Neill H, Perkash I. Skin blood flow changes and tissue deformations produced by cylinderal indentors. J Rehabil Res Dev 1985;22(30):1-6.
  135. Meijer JH, Schut GL, Ribbe MW, Groovaerts HG, Nieuwenhuys R, Reulen JPH, Schneider H. Method for the measurement of susceptibility to decubitus ulcer formation. Med Biol Eng Comput 1989;27:502-6.
  136. Zhang M, Roberts VC. The effect of shear forces externally applied to skin surface on underlying tissues. J Biomed Eng 1993;15(6):451-6.
  137. Zhang M, Turner-Smith AR, Roberts VC. The reaction of skin and soft tissue to shear forces applied externally to the skin surface. J Eng Med 1994;208(4):217-22.
  138. Tam EWC, Mak AFT, Evans JH, Chow YYN. Post occlusive hyperaemic of tissue under static and dynamic loading conditions. Proceedings of 20th Annual Int Conf IEEE/EMBS, Hong Kong; 1998. p. 2294-6.
  139. Herman EC, Knapp CF, Donofrio JC, Sacido R. Skin perfusion responses to surface pressure–induced ischemia: implication for the developing pressure ulcer. J Rehabil Res Dev 1999,36(2):109-20.
  140. Goossens RHM, Zegers R, Hoek van Dijke GA, Snijders CJ. Influence of shear on skin oxygen tension. Clin Physiol 1994;14:111-8.
  141. Krouskoup TA, Reddy NP, Spencer WA, Secor JW. Mechanisms of decubitus ulcer formation--an hypothesis. Med Hypotheses 1978;4(1):37-9.
  142. Bader DL, Gigg SL, Polliack AA. The importance of pressure and time in determining soft tissue status. Proceedings 20th Word Congress of Biomechanics, Amsterdam, The Netherlands; 1994. p. 11.
  143. Ferguson-Pell M, Hagiswa S. Biochemical changes in sweat following prolonged ischemia. J Rehabil Res Dev 1988;25(3):57-62.
  144. Polliack A, Taylor R, Bader D. Analysis of sweat during soft tissue breakdown following pressure ischemia. J Rehabil Res Dev 1993;30(2):250-9.
  145. Pye G, Bowker P. Skin temperature as an indicator of stress in soft tissue. J Eng Med 1976;5(3):58-60.
  146. Goller H, Lewis DW, McLaughuin R. Thermographic studies of human skin subjected to localized pressure. Am J Roentgenol Radium Ther Nucl Med 1971;113:749-54.
  147. Saunders GT, May BJ, Hurd R, Milani J. Lower limb amputations: a guide to rehabilitation. Philadelphia: David Company; 1986.
  148. Mahanty SD, Roemer RB. Thermal response of skin to application of localized pressure. Arch Phys Med Rehabil 1979;60:584-90.
  149. Mahanty SD, Roemer RB. Thermal and circulatory response of tissue to localized pressure application: a mathematical model. Arch Phys Med Rehabil 1980;61:335-40.
  150. Schubert V, Fagrell B. Local skin pressure and its effects on skin microcirculation as evaluated by laser Doppler fluxmetry. Clin Physiol 1989;9:535-45.
  151. Fisher SV, Szymke TE, Apte SY, Kosiak M. Wheelchair cushion effect on skin temperature. Arch Phys Med Rehabil 1978;59:68-72.
  152. Francis JE, Roggli R, Love TJ, Robinson CP. Thermography as a means of blood perfusion measurement. Trans ASME J Biomech Eng 1979;101:246-9.
  153. Patel S, Knapp CF, Donofrio JC, Salcido R. Temperature effects on surface pressure-induced changes in rat skin perfusion: implication in pressure ulcer development. J Rehabil Res Dev 1999;36(3):189-201.
  154. Newman P, Davis NH. Thermography as a predictor of sacral pressure sores. Age and Aging 1981;10:14-8.
  155. Chan AW, MacFarlane IA, Bowsher DR. Contact thermography of painful diabetic neuropathic foot. Diabetes Care 1991;14(10):918-22.
  156. Armstrong DG, Lavery LA, Liswood PJ, Todd WF, Tredwell JA. Infrared dermal thermometry for the high-risk diabetic foot. Phys Ther 1997;77(2):169-75.
  157. Schubert V, Perbeck L, Schubet PA. Skin microcirculatory and thermal changes in elderly subjects with early stage of pressure sores. Clin Physiol 1994;14:1-13.
  158. Naylor PFD. Experimental friction blisters. Br J Dermatol 1955;67:327-42.
  159. Akers CWA. Measurements of friction injuries in man. Am J Indus Med 1985;8:473-81.
  160. Dinsdale SM. Decubitus ulcers: role of pressure and friction in causation. Arch Phys Med Rehabil 1974;55:147-51.
  161. Sanders JE, Garbini JL, Leschen JM, Allen MS, Jorgensen JE. A bidirectional load applicator for the investigation of skin response to mechanical stress. IEEE Trans Biomed Eng 1997;44(4):290-6.
  162. Sanders JE. Thermal response of skin to cyclic pressure and pressure with shear: a technical note. J Rehabil Res Dev 2000;37(5):511-5.
  163. Sewell P, Noroozi S, Vinney J, Andrews S. Development in the trans-tibial prosthetic socket fitting process: a review of past and present research. Prosthet Orthot Int 2000;24:97-107.

Go to TOP.

View Contents

  Last Updated Tuesday, December 2, 2008 12:25 PM

Last revised Thu 5/3/2001