Università di Pisa
Facoltà di Ingegneria
Corso di Laurea Magistrale in Ingegneria Biomedica
[Sottotitolo del documento]
DESIGN OF A SOCKET WITH VARIABLE STIFFNESS
FOR LOWER LIMB PROSTHESIS
Candidato
Relatori
Michele Ibrahimi
Prof.ssa Arianna Menciassi
Ing. Linda Paternò
Prof. Leonardo Ricotti
Index
Abbreviations 3 Motivations 4
Chapter 1: Introduction ... 6
1.1 Incidence and etiology of amputations ... 6
1.2 Human walking ... 7
1.3 Lower limb Prosthesis: a brief history ... 10
1.4 The prosthesis nowadays ... 11
1.4.1 Foot ... 12 1.4.2 Pylon ... 13 1.4.3 Knee ... 14 Chapter 2: Socket ... 17 2.1 Design ... 17 2.2 Suspension system ... 20
2.3 The stump-socket interface ... 25
2.3.1 Residual limb problems ... 25
2.3.2 Overall factors ... 26
2.4 The stump socket interface: biomechanical parameters ... 28
2.4.1 Gait parameters and energy expenditure ... 28
2.4.2 Pistoning ... 29
2.5 The stump-socket interface: physical parameters ... 31
2.5.1 Temperature ... 31
2.5.2 Stresses ... 34
2.5.3 Volume fluctuation ... 39
Chapter 3: Prosthesis Evaluation Questionnaire (PEQ) ... 44
Chapter 4: Pressure ... 48
4.1 Methods ... 48
4.1.1 Textile sensor ... 49
4.1.2 Commercial Pressure sensor ... 51
4.2 Results and discussion ... 54
Chapter 5: Volume ... 63
5.1 Methods ... 63
5.2 Water displacement ... 64
5.3 Water displacement with molds ... 66
5.5 Results and discussion ... 72
Chapter 6: Prototyping ... 78
6.1 Traditional socket realization ... 79
6.2 New methods: Reverse engineering ... 81
6.3 Layer jamming: characterization ... 82
6.3.1 Experimental set-up and results ... 83
6.3.2 Results ... 85
Chapter 7: Conclusions and perspectives ... 87 Bibliography 90
Abbreviations
1. AK………Above Knee 2. AP……….Anterior-Posterior 3. BK………Below Knee 4. CAT-CAM………..Contoured Adducted Trochanteric-Controlled Alignment Method 5. Hi-Fi………..High Fidelity 6. ICS………Ischial Containment Socket 7. IRC………Ischial Ramal Containment 8. KBM………Kondylen-Bein-Muenster 9. MAS……….Marlon Anatomical Socket 10. ML……….Medio-Lateral 11. NSNA………Narrowed Shaped-Narrow Aligned 12. PTB-SC………Patellar Tendon Bearing Supra Condylar 13. PTB-SCSP………Patellar Tendon Bearing Supra Condylar-Supra Patellar 14. PTK……… Patellar Tendon Kegel 15. QUAD……… QUADrilateral 16. SSB……….Specific Surface Bearing 17. SUBIsc……… SUBIschial 18. TF……… TransFemoral 19. TSB………Total Surface Bearing 20. TT………TransTibial 21. VAS……….Vacuum Assisted Suspension 22. VSAP………...Vertical Shock Absorber Pylon 23. TES………. Total Elastic Suspension
Motivations
The design and development of a prosthetic limb is one of the most active and ambitious target in the rehabilitation and prosthetic research, combining the application of engineering principles with medical considerations. The last half century has seen great technological improvements in this field, focusing on the design of fine and advanced components, like electronic knees or ankles or elbows, but it has not encountered the same advancements regarding prosthetic sockets. In upper and lower limb prostheses, the socket is the “Human Machine Interface” (HMI), hence it is the key aspect for the success or failure of a prosthetic device. Indeed, it provides the control of the artificial limb. Even the most brilliant and advanced systems are useless if simplicity and comfort are not guaranteed. A good socket must ensure efficient fitting, load transmission, stability and control by the structural mechanical coupling with the residual limb. It should be designed on the limb anatomy and featured by a modulation of stiffness, in order to minimize shear stresses and normal pressures on tissues. It should be able to guarantee a correct thermal environment and to perfectly fit with the residual tissues, compensating the volume fluctuations over the time.
Today’s sockets are mostly hand-made and realized with a custom design, which cannot accomplish for the residual limb changes nor for different activities and tasks, such as running, walking, ascending the stairs or sitting comfortably. Each activity determines different forces acting on the residual limb and this depends on the socket design, besides the suspension, the knee, the pylon and the foot. The knowledge of the pressure distribution has a key role in order to enhance the comfort. The volume is fundamental too, indeed the residual limb can undergo volume changes in the long and short period, so the stress distribution changes too. On the other hand, stresses are related to the temperature. Skin temperature can increase because of the barrier created by the socket to thermal transfer mechanisms, and this can be related to internal factors, such as the metabolic demand, or to external causes, such as frictional forces. When the temperature increases, perspiration occurs, favoring the socket slippage and the skin problems such as abrasions, skin erythema, ulcers etc. At the same time, slippage is affected by volume fluctuations too, since it alters the fit of the socket. For all these reasons, it is clear that optimizing the interaction between the stump and the socket, is fundamental to guarantee a long-term acceptability.
The aim of this work is to realize a new prosthetic socket, which wants to lay the foundation for a new generation of sockets, perhaps even for upper limbs. A first assessment of the residual limb volume fluctuations and pressure distribution is given, since there is clearly a lack of this information in literature, especially regarding transfemoral amputees. Indeed, two transfemoral amputees have been considered for this thesis. The work will be carried on a bigger sample of amputees, thanks to the MOTU project, funded by INAIL. This will strengthen our conclusions and it will give us the possibility to optimize our design.
According to these parameters the design was realized. The temperature could not be measured on these amputees; however the problem was kept in mind. Indeed, the temperature has been minimized by designing a prototype, made of struts and opened areas, which should let the skin breath. Secondly, the structure wants to accommodate volume variations and to enhance the comfort through the application of the layer jamming principle. Indeed, the layer jamming allows to control the stiffness through vacuum pressure, which is already used in the prosthetic field for the suspension of the residual limb. Hence The strut of layer jamming allows to increase the structure stiffness during specific tasks, while, during the rest of the time, it makes it softer, and it enhances the comfort. Regarding the volume, a manual cabling system has been used to modulate the compression acted by the struts, with the purpose to compensate the volume fluctuations of the residual limb.
Moreover, the socket realization workflow was speed up by 3D printing the components on the base of the the shapes of the amputees’ residual limb. Some rigid struts have been designed, in order to guarantee the amputee’s stability, in case of failure of the layer jamming system. Once done the prototype, tests on bench will follow to confirm the feasibility of this approach and design.
Chapter 1:
Introduction
1.1 Incidence and etiology of amputations
Limb amputations compromise the quality of life of many people, determining physical disabilities. World Health Organization estimates 40 million amputees in the world. Many of them live in developing countries or in places under war, where the absence of a health organization makes harsh to define the amputation causes. However, in places like Afghanistan, Angola, Bosnia Herzegovina, Cambodia, Croatia, Eritrea, Iraq-Kurdistan, Somalia and Sudan, mines represent yet the main cause of amputation for young people and children (70/90 % and 5/15 % of amputees respectively [1]). In developed countries, like USA, approximately 185000 amputations are performed annually, 30 000 of which are at the transfemoral level [2]. In 2005 almost 2 million people were living with a limb loss and, according to statistical projections, this number will double within the 2050 because of the increased life expectancy and the associated risk of vascular diseases and diabetes [3]. Indeed, these are the most common causes of amputation (54%), followed by traumatic events (45%) and cancer (less than 2%) [3]. Chatterjee et al. [4] reported again the vascular diseases as the primary cause, followed by trauma events, cancer and congenital anomalies.
World Health Organization estimates 350 million of diabetics in the world and this leads to increase the possibility of vascular diseases, which can lead to limb amputations as evidenced in developed countries. As reported by the Amputee Coalition in the USA, ‘nearly half of the individuals who have an amputation due to vascular disease will die within 5 years’ [6]. This mortality rate is higher than the five year mortality for breast cancer, colon cancer, and prostate cancer [7]. Besides, more than the 55% of the people who underwent a leg amputation will have the other leg amputated in 2‐3 years. Thinking about Italy, 3 million of people live with diabetes and this increases the probability of vascular disease. The Italian Ministry of Health reports nearly 10000 new amputations per year, 3000 of which are transfemoral amputees.
According to these data, amputations are exponentially increasing, affecting people’s life more than commonly expected. The importance of an efficient and comfortable prosthesis is evident in order to recreate a functional gait cycle. The following section describes the human walking, evidencing the complex biomechanics which should be reproduced by a prosthesis.
1.2 Human walking
Lower limb amputation compromises the daily routine of people. Indeed, although the human walking is one of the first things learnt at the early stages of life, its apparent simplicity hides a difficult understanding, which amputees struggles to learn again.
The gait cycle is made of a repetitive sequence, starting when the heel hit the ground (e.g. heel strike) and ending with the contact of the same foot with the ground. This motion guarantees the lowest energy expenditure at the natural gait speed. Amputees have to learn a new way of walking, which includes the prosthesis.
The lower limb consists of four major parts, namely a girdle formed by the hip bones, the thigh, the crus and the foot. It is specialized for the support of weight, adaptation to gravity and locomotion. The major bones are the ilium, ischium and pubis, in the hip region, the femur, tibia and fibula, in the lower region (Figure 1). It has to be noticed that most of the leg skeleton has bony prominences and margins, such as the anterior superior iliac spine, the greater trochanter, the superior margin of the medial condyle of tibia, and the medial malleolus.
If these areas are pressed against rigid walls, as it happens in prosthetic sockets, it can cause pain.
The leg motion during the gait cycle involves each foot in turn, advancing forward as a step or supporting and balancing the body weight during the advancement of the contralateral leg. It can be described by time segments, each characterized by an event as showed in Figure 2
.
Two major phases can be distinguished: the stance phase, when the foot is in contact with the ground (60% of the gait cycle), and the swing phase, when the contralateral leg swings. This time pattern is possible thanks to the activity of the lower and upper limb muscles, which must overcome gravity and guarantee a forward propulsion. The major muscles involved in the human walking are the ileoipsoas, gluteus minimus, medius and maximus, tensor fascia latae, quadriceps femoris, tibialis anterior, triceps surae and hamstrings.The stance phase begins with the period of weight-acceptance, specifically with a double-limb stance, where the body weight is transferred from one leg to the other. It is characterized by a leg deceleration, since the foot has to stop after having swung. This action involves simultaneously the ankle, knee and hip muscles recruitment, while the contralateral leg is initiating the swing phase (acceleration phase).
Figure 2: Time segments of human walking. Image adapted by [7]
Considering the gait cycle described in Figure 2
,
at the Initial Contact, the ankle dorsiflexion is followed by the plantar flexion, caused by the eccentric contraction of the tibialis anterior as the foot flattens on the ground (Loading Response). Simultaneously the knee is flexed under the eccentric control of the quadriceps. The hip is flexed too and it begins to extend on the pelvis, as the trunk keeps moving forward. The forward motion is controlled by contraction of the hipextensors, the gluteus maximus and the long hamstrings. The left foot leaves the ground, while the right hip continues to rotate internally, entering the single-limb support time of the stance phase. The right ankle begins to passively dorsiflex as the tibia leans forward and the pelvis lowers on the left and moves laterally to the right. In this way the Centre Of Mass (COM) of the body is aligned over the right foot in order to allow balance. During this period, the pelvis and COM of the body are rising. In the middle of single-limb stance phase, the gait cycle ends its decelerating phase and begins the acceleration phase as the left heel rises. As the knee extends, the pelvis maximally rotates externally and the right hip maximally extends in preparation for the opposite heel strike. While the left leg accelerates, the weight-release phase begins on the right leg. Iliacus, psoas, and tensor fascia latae muscles produce the right hip flexion, and the left hip internal rotators cause the forward rotation of the right pelvis, contributing to the horizontal forces of the forward propulsion of the same. When the left toe finally leaves the ground, the swing phase is beginning and a step occurs. In the middle of the swing, the toe reaches a minimum height of less than 2.5 cm. This allows the minimum energy expenditure, thanks to a low work made to elevate the limb [8]. The second half of the swing phase is characterized by a decelerating phase of the right leg, as the left heel strike is preparing.
Regarding the gait cycle of amputees, it is obviously altered and, depending on the prosthesis goodness, different problems can be observed. Some examples are lateral trunk bending, toe hitting ground, wide walking base, vaulting, uncorrected knee flexion etc. [9].These depend on the prosthesis components and their assembling, highlighting again the importance of an efficient prosthesis and a good mechanical coupling with the stump.
1.3 Lower limb Prosthesis: a brief history
Prostheses have a long history, looking back to the Ancient Egypt. The most ancient prosthesis dates back to the period 950-710 BC (Figure 3 a). It is an artificial toe of a female mummy, found near Luxor and made from leather and molded and stained wood. The Lancet toe is another example of prosthesis dating back to 600 BC and with cosmetic purposes (Figure 3 b).
Figure 3: (a) The Cairo toe, (b) the Lancet toe, (c) Roman leg, (d) Gotz hand, (e) Parmelee and (f) modern prosthetic leg
A prosthetic bronze leg was discovered in an ancient tomb in Capua (Rome, 1910). It is dated back to 300 BC (Figure 3 c). The leg was made to replace the leg below the knee. It was kept at the Royal College of Surgeons but it was destroyed during World War II.
After that, the Renaissance was a rebirth period also for the prosthetic field too. Prostheses began to be made of iron, steel, copper and wood.The first technologically advanced iron hands, lately found, belonged to a German mercenary, Gotz von Berlichingen (1508) (Figure 3 d). These hands were suspended by leather straps and controlled by a series of springs.
In the 16th century, Ambroise Parè introduced modern amputation surgery techniques and new
prosthetic designs. However, the first advanced prosthetic limb is dated 1863. It was an artificial leg, patented by Dubois D. Parmelee, with a suction socket, polycentric knee and multi-articulated foot (Figure 3 e).
In the 20th century technological advancement and a great attention towards the casualties of the
Great Wars led to the development of new artificial legs, focused on the design of custom sockets , artificial electronic knees and feet. In this way, modern prosthetic leg is introduced (Figure 3 f).
1.4 The prosthesis nowadays
Lower limb prostheses can be classified according to the level of amputation, that is partial foot amputation, ankle disarticulation, transtibial amputation, knee disarticulation, transfemoral amputation, hip disarticulation and hemipelvectomy (ISO CODE 99999) [10]. Generally, long residual limbs allow better stability and control of the prosthesis. This way, muscles can be saved, avoiding permanent abducted and flexed behaviors, especially in transfemoral amputees. However, transtibial stumps should not be longer than 12 cm in case of arterial insufficiency, else the healing would be difficult [11].
Otherwise, prostheses can be classified according to their structure and components. Exoskeletal prostheses are characterized by an external structure, which has the bear loading function. They are mainly used for cosmetic purposes, as bath prosthesis, or as work prosthesis, wherever there is powder or dust. This kind of prosthesis is passive (Figure 4 a), because there is no energy supply helping the motion. The second type of prosthesis is modular, which can be semiactive, in the case of systems able to store and release the energy (Figure 4 b), or robotic, if the ankle or/and the knee are actively controlled by a microcontroller (Figure 4c).
Besides these systems, new treatments for osseointegrated prostheses have been introduced in 1990 in Sweden and new studies (Figure 4d) have been recently carried on in United Kingdom, Australia, Hungary and France [12]–[14]. However, these new kinds of prostheses need to be deeply evaluated for the high risk of infection and bone fractures (Figure 4 d).
Figure 4: Cosmetic wood leg worn by Aimee Mullins at the 1999 fashion week (a), Cheetah legs made popular thanks to Oscar Pistorius (b), Robotic legs worn by Huge Herr (c) and an osseointegrated transtibial leg (d)
To date, modular prostheses constitute the standard of care, thanks to their versatility. Indeed, they are made of a socket with a suspension system, a knee if needed, a pylon and a foot, which may be changed when damaged, or for other reasons.
After the surgical procedures, a temporary prosthesis is assigned to the amputee. During the rehabilitation period, the residual limb undergoes a shrinking and rectifications of the socket can be made. After 6/18 months the permanent prosthesis is given. Each component is aligned in order to reproduce the functional line passing through the epicondyles of the leg bones. This is made in order to optimize the Range Of Motion (ROM) (e.g. the movement ability of the leg joints
confined by the socket) of daily movements and gait cycle [15].
The next sections describe all the modular prosthesis components, focusing later on the main topic: the prosthetic socket.
1.4.1 Foot
As each prosthetic component, the foot is chosen by the prosthetist according to the patient’s weight, activity level, and stump features, like length or muscle tonicity. Wrong components can increase the metabolic cost, wrongly activate the muscles, and decrease the gait symmetry. For this reason, many studies have been done in the last 20 years in order to understand the mechanic behavior of prosthetic feet and their biomechanical implications [16].
Prosthetic feet can be mainly classified in non-articulating or articulating [17]. These can be classified into monoaxial or pluriaxial feet. The non-articulating feet or SACH (Solid Ankle-Cushion Heel) are light and cheap solutions (Figure 5 a). They have a wooden keel, covered by a rubber material, which reproduce the foot shape. The heel stiffness can be changed by varying the heel material in order to enhance the cushioning effect. In the ‘80s they were the standard of care, until the development of articulating feet.
The monoaxial articulating feet allow the flexion on the planar and dorsal planes, whereas non-articulating feet do not allow any movement. The pluriaxial non-articulating systems give greater Degrees of Freedom (DoFs), since they allow movements in the sagittal, coronal and transverse planes at the tibiotarsal level. However, these systems weight more and are not suitable for short transfemoral residual limbs or for transtibial amputees with unstable knees.
For this purpose, dynamic articulating feet have been designed. They can be pluri or monoaxial articulated. The Seattle foot has been developed as the first dynamic socket in the ‘80s [18]. The concept is to have a foot, whose design and material allow to store the energy of the shocks, by spring like elements, and to release it as needed. This lowers the energy expenditure and improves running tasks. Other systems developed over the years are the Carbon Copy (Ohio Willow Wood
Company Inc., USA) or the more popular Flex-Foot/ Flex Foot Cheetah (Ossur, Iceland) (Figure 5 b, c). Moreover, multiaxis ankle motion can be achieved with the prosthetic feet thanks to multiaxial ankle components, like the Impulse Ankle (Ohio Willow Wood Company, Inc., USA), which features inversion/eversion and axial rotation. This is appropriate for active individuals and for any amputee who needs improved ankle motion, for example to accommodate gait on uneven terrain.
Figure 5: (a) SACH foot, (b) FlexFoot, (c) Flex foot Cheetah and (d) the Reflex foot (Vertical shock absorbing pylon, dynamic response, multiaxis, and transverse motion)
1.4.2 Pylon
The pylon is the element of connection between the foot and the prosthetic knee or the foot and the socket. It can be made of stainless steel, aluminum, carbon or titanium, according to the needed mechanical features. Classic pylons were rigid elements, whereas nowadays they can incorporate compression elements too. These systems
,
called Vertical Shock Absorbing Pylons (VSAPs) (Figure 6), aim to absorb the cyclic loads applied on tissues by the ground reaction, lowering the risk of pressure blisters and abrasions. Their mechanical responses to specific loads have to be better evaluated, as their effect on amputees’ gait. No alterations of the gait seem evident, and a reduction of the energy cost has been proven [16], [17].Between the pylon and the socket (or the feet), torsional rotator adaptors can be fixed. There are two types of rotator adaptors, one for transfemoral amputees and one for both transfemoral and transtibial ones. The first type is placed at the base of the socket and it is used to rotate all the assembly in order to facilitate simple tasks like wearing a shoe (Figure 6). The other type can be placed between the pylon and the foot or at the base of the socket and it is used to lower the shear stresses between the socket and the skin, by allowing micro rotation of the components in the transversal plane [19]. Indeed, when the upper body rotates in the transverse plane and the feet are firm on the ground, a distortion of the tissues can occur. To date rotator adaptors are only
adjustable by a prosthetist, who is responsible of setting the stiffness, not adaptable to the different activities of daily life. A variable stiffness rotator has been developed by Pew and Klute [19], demonstrating a reduction of moments on tissues in the transverse plane, while maintaining a good stability at different walking speeds. Loads were monitored by simply applying a load cell beneath the adaptor. Some research efforts have been made towards the development of ad hoc systems, positioned between the socket and the pylon and able to monitor the loads transmitted by the ground to the socket [20]. This is extremely useful to segment the gait phases of amputees in real time and it eventually allows to understand the relation between the ground reaction and the stresses distribution at the socket interface.
Figure 6: (a) Shock absorber pylons reported by Klute et al. [16]; from left to right: Pyramid Pylon (Blatchford Endolite, Chas A. Blatchford & Sons Ltd., Hampshire, U.K.), ICON Shock Pylon (Flex-Foot Inc., Aliso Viejo, CA) and Total Shock Pylon (Century XXII Innovations Inc.,Jackson, MI). (b) Variable stiffness torsional adapter (adapted by [19]) and (c) rotation adapter in a TF amputee (adapted by [17]).
1.4.3 Knee
The knee is a fundamental component of a prosthesis since it has to allow good stability in stance phase and to guide the swing of the artificial leg during the dynamic phase. A good knee should optimize the energy expenditure, absorbing the vertical shocks (such as at the heel strike). In addition it should be able to shorten the prosthesis in order to avoid the tiptoe hitting over the ground during the swing phase (clearance of the toe) and allowing the knee to bend during tasks like sitting.
Three main phases has to be considered when it comes to knees: Heel Strike, Mid-Stance and Toe Off (Figure 7). During Heel Strike the reaction force is not aligned with the load axes. The knee center of rotation is posterior to the line, leading to an extension momentum, which avoids the knee flexion. Reaching Mid-Stance, the reaction force and the knee center of rotation are aligned along the load axis until the knee center of rotation becomes anterior to the load axis. This
initiates a flexion momentum, bending the knee (Toe Off). A knee center of rotation too posterior increases the effort to flex the knee and a knee center of rotation too anterior flexes the knee too early.
Prosthetic knees can be classified into monoaxial or polycentric (fix or moveable centroid respectively) and by the mechanism which controls the speed and the ease of the swing. The monoaxial knees can have mechanical hinges controlled manually by constant or weight activated friction. Both monoaxial and polycentric knees can be controlled manually by hydraulic or pneumatic systems. Otherwise they can be electronic [17].
Figure 7: (a) Rapresentation of a static non-amputated trochanter, knee, ankle alignment, (b) phases of the static controll of a monoaxial knee and (c) sagittal static allignment (femur has to be adducted 5°). Images adapted modified by [21]
Manually controlled knees are manually blocked by a lace, along all the gait phases. It is a system often advised for geriatric users and people with low stability in stance phase and with the high risk of fall, since the knee maintains extended the leg. Friction knees are based on a mechanical system, whose friction can be set by the prosthetist. Otherwise, they can be activated by weight bearing, during the stance phase. This way, the stability is not optimal, hence they can be used for light persons. The other systems can be polycentric too, i.e. the centroid translates along the swing phase (maintaining the previous considerations) and increases the toe clearance by shortening the shank with the articulated quadrilateral system.
Monoaxial or polycentric knees can be controlled by a pneumatic or hydraulic systems, which have the main advantage to allow the users change their speed. Indeed, they can be prescribed for active patients too, but they require maintenance and they are heavy. Microprocessor knees (C-leg, Rheo knee etc.) are advanced systems with an on-board computer and load sensors,
accelerometers, gyroscopes, and joint angle sensors, which detect the gait phase and speed in order to improve the static and dynamic phases. However these systems are expensive and heavy [22].
The prosthesis alignment must be optimized considering each component in the lateral and sagittal plane, according to the physiological lines as showed in Figure 7. The first alignment is done on bench, made straight after the prosthesis realization. Then, it has to be evaluated in static and dynamic conditions on each patient. It is extremely important because it affects the knee function and the overall gait biomechanics [9]. Besides that, it influences the stresses distribution at the socket-stump interface. Limburg et al. [23] studied the effect of ankle joint realignment on the pressure distribution of a transtibial amputee, showing a significant reduction of the average pressure at the subpatellar (-40% of body weight per cm²) and tibia-end (-30%) areas in stance phase. Sanders et al. [24] investigated different socket-shank alignment and verified that alignment changes had a localized effect on interface stresses (pressure and shear stresses). However, the effects of the alignment and previous described components have not been deeper investigated in this research as done for socket designs and suspension systems, which have been the focus of the research.
Chapter 2:
Socket
2.1 Design
According to the amputees’ residual limb features, such as stump dimensions, age, lifestyle and activity level, the socket is shaped by the prosthetist, keeping in mind what kind of suspension will be used. These systems are indeed chosen together, since they influence each other. However, the socket shape can be classified in different main designs, for each level of amputation (Figure 8).
Figure 8: Diagram of the socket designs in lower limb amputees.
Transtibial sockets
In the early ‘50s the Patellar Tendon Bearing (PTB) socket was the standard for transtibial amputees. Its name is related to the loads distribution inside the socket, focused on tolerant pressure areas (Specific Surface Bearing socket (SSB)), like the patella [8]. In this way, bony prominences are relieved. The special trimlines allow a grip on the femoral condyles, and this is enhanced in the PTB SupraCondylar (PTB-SC) version and in the SupraCondylar/ SupraPatellar (PTB-SCSP) one. These designs are characterized by higher lateral and lateral/ anterior walls respectively [17]. The Patellar Tendon Kegel (PTK), also called Kondylen-Bein-Muenster (KBM) socket, is another design addressing the pressure on specific areas (SSB), where the proximal walls (Figure 9) are extended and there is the maximum grip on the condyles. These configurations are
taken into account when there is a lack of lateral knee stability. In this way there is a reduction of loaded anatomical areas, but this increases the pressure on specific regions, and several skin problems can occur. For this reason, in the ‘80s Total Surface Bearing (TSB) (Figure 9) sockets were introduced. The main concept is to guarantee a more uniform stress distribution by applying the pressure on wider areas [25].
Figure 9: (a) Specific surface bearing sockets and (b) total surface bearing socket. TSB sockets have trimlines similar to PTB sockets, but they are realized in order to completely adhere to the skin, even if it is not evident from the figure.
Transfemoral socket
Regarding transfemoral amputees, in the ‘50s the most diffuse socket design was the quadrilateral socket (QUAD) (Figure 10). As PTB sockets are focused on bearing the loads at the patella area, the QUAD sockets are based on the idea of an ischial weight bearing system. Its name is due to the quadrilateral shape in the transversal plane. Quadrilateral sockets tend to compress the muscles in the antero-posterior direction. Therefore Radcliffe et al. [26] highlighted problems like the proximal shifting in the medio-lateral direction, and femur hitting at the distal lateral wall, caused by pelvis’ rotating action towards the medial side. Besides, residual limbs have a fixed abduction angle caused by the amputation, which QUAD sockets cannot correct, as confirmed by Long et al. [27]. This problem can be overcome by an optimal alignment (Long’s line). In order to avoid the lateral gapping and the discomfort caused by quadrilateral sockets, Ischial Containment Sockets (ICS) were investigated in the ‘80s. This design is characterized by a higher medial wall, which grabs the ischial tuberosity. Some variations are: the Normal Shape-Normal Alignment (NSNA) [27], the narrow Medio-Lateral (ML) [28], the Contoured Adducted Trochanteric-Controlled Alignment Method (CAT-CAM) [29] and the Marlon Anatomical Socket (MAS) [25]. All of these are focused on the improvement of the femoral alignment within the prosthesis, in order to enhance the lateral stability (Figure 9). For this reason, they are characterized by extended
medio-lateral trimlines, which contain the ischial tuberosity and ramus. The CAT-CAM has been modified over the years and it led to the design of the Sabolich socket (Figure 10). This kind of socket is made of an external structure, made of rigid struts, and an inner socket made of polyethylene or silicone elastomer. Some other similar examples are the IPOS, the ISNY (Icelandic Swedish New York) and the SFS (Scandinavian Flexible Socket), which differ for the frame fabrication techniques and the material used in the inner socket. Recently the Hi-fi socket has been patented [30]. It is a socket made of struts, which compress the tissues in specific areas and let them escape in others, but it will be better discussed in the Stresses section 2.5.2. The MAS design is another design developed in the 1999, which claims to enhance the comfort and ROM by lowering the posterior trimlines at the gluteal area. The main drawback is the greater compression applied in order to realize such socket. In the last years, the subischial concept has been suggested again [31], [32], also thanks to the chance to use vacuum-assisted suspensions. The new design is the Sub-ischiatic Northwestern socket (SUBIsc), which provides a new solution to medium-long thighs, aiming to enhance stability, gait and comfort [31]. It has lower trimlines and allows a wider ROM of the hip, contrarily to the ICS sockets (Figure 10
).
It can be made thanks to the use of a vacuum suspension, which enhances the stability and allows to lower the trimlines [33].Figure 10: (a) Design of the possible socket shapes in a frontal view and (b) different frame designs of flexible sockets
2.2 Suspension system
Suspension systems can be classified into harness and suction systems for both above and below knee amputations (Figure 11) [8], [21], [33]–[35]. Harness systems are characterized by cabling systems, possibly paddles and corsets. They are the first systems developed [36]–[38] and, nowadays, are still used to enhance the stability of short limbs, combined with other suspension types. On the other side, suction suspensions are characterized by subatmospheric pressure inside the socket. This can be achieved by pulling away the air through an unidirectional valve at the bottom of the socket, and can be applied directly on the skin or protecting the stump with a roll-on liner. Roll-roll-on liners are soft covers, made with silicroll-one or similar elastomeric materials, which can be locked to the socket by distal mechanisms, e.g. pin-lock, magnetic lock, lanyard lock. They can include hypobaric seals, in order to enhance the subatmospheric environment, or can be used with a vacuum pump (Vacuum Assisted Suspension (VAS)), able to guarantee negative pressure during all phase of the gait cycle.
Figure 11: Classification of the suspension systems. The sleeve is an optional component, but usually advised. The textile fabrics configuration has been recently adopted by Martin Bionics.
Each system has its own drawbacks and benefits, and it is carefully chosen by the prosthetist on the base of patient and stump features (activity, age, health problems, stump length, muscular tonicity etc.).
Harness Suspension System
The most used transfemoral harness suspensions are the Total Elastic Suspension (TES), and the rigid Silesian belt, showed in Figure 12. The TES is a neoprene belt, which surrounds the waist above the iliac crests and the proximal regions of the socket in order to provide suspension. On the other hand, Silesian belts attach to the anterior and lateral areas of the proximal prosthetic socket and pass over the opposite iliac crest. These systems can be adopted as primary suspension systems or as an auxiliary to an additional one. They are simple to don and suitable for low activity patients. The main disadvantages are related to frequent pistoning movements (axial movement of residual limb tissues and bone respect to the socket), reduced comfort, heat and dermatitis. Other auxiliary rigid systems are the suspenders, the pelvic band and hip joint. This last option uses a single axis hip joint positioned on the lateral socket wall and connected to a pelvic band, enclosing the iliac crest. The side joint and band limit the rotation of the prosthesis and stabilize the system.
Figure 12: Belts, suspenders and straps in various configurations used for TF (a) and TT (b) amputees, adapted by [17]
Regarding transtibial suspensions, a commonly used elastic system is the supracondylar cuff strap. All of the possible variants consist of a multipart strap that attaches to the sidewalls of the socket and encircles the distal thigh. This strap is easily applied, cheap and comfortable. Hence, they are good for low activity patients and when impaired hand function limits the grip coordination. Waist belts are rarely used. If adopted, they are usually connected to a high corset around the thigh by
side joints applied at the socket wall. This solution could be suggested in case of a high unstable knee, but it is often avoided because it leads to muscle atrophy.
Otherwise, the harness system can be embedded in the socket as well, which means a system of cables or straps manually adjustable by the patient [39], [40]. These systems are mechanism of volume control as well. The sockets provide the suspension with the aid of liners or a textile fabric. A suction system can be added too by creating a distal ‘cup’, as done in some available solutions of the infinity socket introduced later [41]. Textile fabrics have not been much used until now because of their high friction, but a recent development of Martin Bionics Inc. created a new paradigm. They created a ‘blossom’ like structure, which wraps around the distal residual limb and it is then covered by a textile fabric structure (Air Hammock suspension), studied for other applications too [42]. However, no clinical evaluations have been found in the literature.
Suction Suspension System
Suction suspensions are characterized by a sub-atmospheric environment inside the socket, achievable by applying the vacuum through different ways. The residual limb can be pushed into the socket by simply applying a lotion on the skin (skin fit suction) and the air is expelled through a unidirectional valve. To date this option is often not recommended because of the stump high sensitivity. In the ‘80s the concept of liner started to take place, as reported at the 1984 AOPA Assembly by Ossur Kristinsson [43], inventor of the Iceross liner and founder of the Ossur company. Liners consist of a stretchable cylinder closed at one end, which can be made of silicone elastomer or gel, poly-urethane, thermoplastic elastomer, mineral oil gel or closed cell foam materials. The distal part has a thickness of about 6-7 mm, and lateral walls of 2-4 mm. This ‘sock-like’ system is rolled onto the residual limb, transferring the friction action to the liner-socket interface. The distal end provides cushioning and a better mechanical loads transfer as reviewed by Klute et al. [44], while its general compression redistributes evenly the loads. Their main drawback is the low thermal conductivity. Liners are thermal insulators [45] and sweat, moisture and unpleasant odors can occur, besides the slippage of the liner itself.
Liners with distal lock can be with a strap, magnetic mechanism or pin lock (Figure 13). The first type has a lanyard strap which passes through a hole across the socket and it is tighten and attached to a velcro-strap on the outer socket. It is suitable for geriatric users for its simplicity, but it does not stabilize the stump by rotational movement. Liners with a pin lock mechanism offer a greater stability, but the main drawback is the “milking effect” (excessive elongation of distal
tissues) [46]. Liners with a magnetic lock are another solution (Otto Bock Dynamic Vacuum). Eshragi et al. [47] proposed a questionnaire to the patients, who declared to prefer the magnetic suspension due to ease of doffing and reduced noise, notwithstanding the larger pistoning. Suction solutions (Figure 13 a) are simple liners with a cushioning end. By donning the prosthesis, the air is expelled through an unidirectional valve or a vacuum assisted suspension (VAS) system, which imply the use of a vacuum pump. Introduced in the prosthetic field by Caspers (1995), VAS can be mechanical or electrical (Ottobock Harmony P3, Harmony e-Pulse, Ohio Willow Wood VS, Ossur Unity).
Figure 13: (a) Suction suspensions, consisting in a suction solution with a basic liner and a sleeve (an internal additional sleeve can be added) and a Seal-In liner. (b) Liners with different locking mechanisms
Mechanical pumps create the vacuum after a cycle of 30 steps, instead electrical systems need a dozen of steps, but they last as long as the battery allows it. A hybrid solution has been developed by Fatone et al. [33], using an electrical system to create the vacuum and a mechanical coupling to maintain it along the walk. VAS enhances the total contact of the stump-socket, reduces the pistoning and increases the control in swing and stance [62], [81]. They also seem to improve wound healing. Interesting fact is that they cause a volume gain instead of a volume loss, because of lower negative pressures during the swing phase and positive pressure during stance, compared to usual suction sockets [51]–[53]. The counterpart is that VAS systems need maintenance, some patients complain of difficult donning [48] and the socket weight increases. When using VAS or unidirectional valves, the use of a sleeve is usually recommended to avoid air leaks, although this solution is not feasible for transfemoral amputees, since there is less space to roll over the sleeve. Therefore, Seal-In liners have been developed to replace sleeves. They consist of one or more adjustable silicone bands, which wrap the liners at mid-length and avoid pressure losses. A survey
on 90 transfemoral amputees stated a more significant satisfaction for Seal-in liners during fitting, sitting, assessing significantly more problems for common suction in terms of sweating, wounds, pain, irritation, pistoning, swelling, smell, and sounds [35]. However they are usually not preferred for geriatric users, because of their uneasy donning/ doffing [54], [55].
2.3 The stump-socket interface
2.3.1 Residual limb problems
Amputations can occur for different causes, changing the patient life. The post-operative phase is one of the most delicate phases, for the stabbing pains on the residual limb, phantom pains and edemas [8]. Overcome the first weeks, a temporary prosthesis is realized and after some months, a definitive prosthesis can be given, depending on the patient’s status. Amputees have to learn a new way of walking through a rehabilitation process, which can last a few years. One of the hardest and most important aspect for a prosthetic limb, is the fitting of the socket [56]. Indeed, the stump-socket interface defines the distribution of forces and pressures acting on the residual tissues of the stump [57]–[59], influencing the patient comfort. High stresses applied on the skin can lead to pressure ulcers, sensitive skin, dermatitis and hyperemia and vascular problems, especially if distributed unevenly (Figure 14) [59]. This modifies the blood perfusion in the soft tissues [25], [60]–[62], causing a temperature increase. As a consequence, a moist environment is favored and tissue deformations can occur [63], increasing the risk of injuries and tissue breakdowns [64], [65]. In the worst scenario, these skin problems can evolve in chronic infections, requiring a reamputation [66]. Hence, besides tissue injuries, stress distribution at the interface between the limb and the prosthesis can be the cause of many discomforts. High shear stresses can worsen the stump thermal conditions too, already altered by the low thermal conductivity of the socket [62], [63], [67].
It is known that stumps are characterized by a change in their volume and shape over time, in the long and short period [68]. These volume fluctuations affect the prosthesis stability by reducing the stump socket fit. Considering an increased volume, a pressure increment occurs in turn at the stump-socket interface. However, it is known that tissues necrosis might occur if pressures higher than 8 kPa are applied for long periods [69]. On the other hand, a lower volume causes abrasive relative movements, especially in the axial direction (pistoning effect) [25], [64], [65], [70].
It obviously comes that lot of amputees reject the prosthesis or have a low level of satisfaction, mainly related to poor comfort, function and control over the socket [71]–[75]. This fact can induce an overuse of the unaffected limb with following problems such as tenosynovitis, epicondylitis, and tissue damages [76], [77].
Figure 14: A diabetic foot which will have to be amputated for necrosis (a), ulcers on a transtibial stump (b) and evolution of a bad stresses distribution (c) and.
2.3.2 Overall factors
Although all the prosthetic components are extremely important to obtain a functional gait, the most important impact is related to the socket, since it is the link to the human subject. To avoid discomfort and skin pathologies, which can lead to the abandonment of the prosthesis, the realization of a socket must consider a series of parameters.
The residual limb volume can greatly vary during the first year of amputation (early post-operative period) and after that, it undergoes a steady state. However daily changes keep happening, depending on the daily routine, activity level, diet and age. The loss of fluids, meant as perspiration of tissues, is a common daily phenomenon, which becomes greater in presence of the socket. The knowledge of these parameters is related to the physiology of an amputee, so it should address the socket design, i.e. materials (thermal conductivity, stiffness etc.), suspension and shape (compliant systems, fenestrations etc.) (Figure 15).
Figure 15: Representation of the physical parameters which each suspension, shape and material of a socket should optimize, and the external overall factors affected by these parameters
On the other side, these factors influence the residual limb load distributions (pressures and shear stresses). Volume gains enlarge pressures and reduce the pistoning, while a temperature increase creates moisture and slippage, so higher shear stresses and pistoning. Shear stresses, slippage and pistoning are related to the coefficient of friction. If the coefficient of friction is too low, the skin-socket interface, liner-skin-socket or skin-liner can result in slippage, causing sores, blisters and an increase in body temperature, which can lead to perspiration and slippage. As a consequence high pressures are needed to bear the loads [78], [79], increasing ischemic risks. If the coefficient of friction is too high, this can lead to high shear stresses, tissue distortion during donning/ doffing and ambulating, with the risk of skin breakdowns.
All these factors are related to each other and they affect the overall comfort and walking biomechanics (Figure 15), besides the influence of components alignment and the prosthetic components themselves. The next paragraphs describe the research efforts to evaluate these parameters.
2.4
The stump socket interface: biomechanical parameters
2.4.1 Gait parameters and energy expenditure
The goodness of a prosthesis can be evaluated by measuring external parameters like gait parameters and energy expenditure. Indeed, these can be compared to healthy subjects in order to reach the best performances. Given the same components, such as knee, pylon, foot and an optimal alignment (based on patients’ specific issues), these values can be a great mean to evaluate socket and suspension on an amputee. Safari and Meyer [57] reviewed the last 15 years of works, involving quantitative data of the effects of sockets and different suspensions. Among 15 works, 11 evaluated gait parameters and only four regarding a comparison between the designs and suspensions. It was stated that TSB sockets improved gait symmetry and further symmetry was achieved with VAS. VAS systems has the advantage to allow a better control of the residual limb volume [80]. Only one work involved the evaluation of the energy cost, which was reduced for TSB sockets with a Seal-In liner, compared to a liner-sleeve suction socket.
For a non-amputee, energy cost increases linearly with walking speed (aerobic metabolism), and anaerobic mechanisms are involved after a threshold value. For amputees, the threshold is reached before than healthy subjects. Metabolic cost increases as +15% for Syme amputees, +25% for traumatic TT, +68% for traumatic TF, +40% vascular TT, +100% vascular TF. Since the ‘70s, energy cost of walking has been evaluated on amputees with QUAD sockets, reporting lower gait speeds than non-amputees, in order to maintain minimal energy expenditure [81]. Flandry et al. [82] compared QUAD and ICS sockets, showing 56% less energy expenditure in the latter, while Gailey et al. [83] showed 20% less energy expenditure. Despite this, no significant difference in the heart rate was observed between the two designs, at two different speeds. Oxygen consumption at slow speeds was neither significantly different for the two designs.
As previously described, in the last decade new TF designs have been proposed, trying to enhance stability, gait and comfort, such as the Northwestern and Hi-Fi sockets. While the SUBIsc socket uses VAS system to gain stability, the Hi-Fi socket is made of a frame of 3-4 struts, extending longitudinally, which compress the limb [32] (Figure 18c). These designs have been the center of investigations of recent works, which tried to understand the brim removal effects. Indeed, the traditional sockets usually cause discomfort at the perineal area and this can lead to abducted thighs. Early studies on brimless sockets (SUBIsc) showed gait parameters at least comparable or even more symmetrical than ICS and a preference towards brimless was stated [57], [49], [80],
[84]. Kahle et al. [85] compared ICS to Hi-Fi sockets using a 2 minute walk test on 13 patients. The mean distance walked was 91.8 m ± 22.0 for the ICS and 110.4 m for the Hi-Fi socket. The amputees did compile a self-report too, assessing their increased balance confidence during daily activities with the Hi-Fi socket [49], [85]. Alley et al. [86] reported an improvement in cadence and step length as well for a transfemoral amputee.
2.4.2 Pistoning
Transtibial
Reference Socket and Suspension Pistoning (mm)
Board et al. 2001 [87] TSB with VAS 1(liner) and 5(tibia) TSB with valve 33(liner) and 44(tibia)
Klute et al. 2011 [88] TSB with VAS 1 ± 3
TSB with pin lock liner 6 ± 4 Gholizadeh et al. 2012 [54] TSB with liner 5.4 ± 0.6 (gait)
TSB with Seal-In liner 2.5 ± 0.4 (gait) Eshraghi et al. 2012 [47] TSB with liner 5.8 ± 0.8
TSB with Seal-In liner 2.8 ± 0.5 Brunelli et al 2013. [89] TSB with liner 12.4 ± 5.6
TSB with Seal-In liner 6.1 ± 3.1
Transfemoral
Convery et al. 2000 [90] QUAD and liner 40 ± 0.5 Kahle et al. 2013 [50]
SUBIsc with liner and
VAS 14 ± 8 (gait)
ICS with liner and VAS 25 ± 9 (gait)
Table 1: Recent works about pistoning. If ‘gait’ is not specified, static loadings are applied. Safari and Meier [46] reports a more detailed description about transtibial pistoning.
Suspension system efficiency is evaluated by measuring the pistoning, term used to address the vertical movement of the stump inside the socket. It can be relative to the skin-socket interface
and bone-socket or liner-socket interface. Measurement techniques involve radiography, ultrasound, computerized tomography, photographic techniques and system motion analysis with markers [35], [91]. As with pressures and shear stresses, the pistoning can be measured in static conditions or in dynamic conditions. A static evaluation can be made by an amputee standing on two legs, standing on the prosthetic leg or by applying a load on the amputee’s prosthetic leg. Otherwise, the dynamic pistoning can be evaluated during the gait or by applying repetitively external loads on the artificial leg.
The works of Cluitmans et al. [118], Datta et al. [92] and Hachisuka et al. [93] reported an improvement of users’ ‘pistoning satisfaction’ between PTB sockets and TSB sockets, as confirmed by Yiǧiter et al. [94]. Recent works focused on different liners. Dermo liners (suction sockets) and Seal-In liners have been investigated, reporting more satisfaction for transtibial amputees [95]. Table 1 reports some recent works assessing a pistoning improvement when using Seal-In liners. About VAS, only a few works reported objective observations, but an improvement is still evident. The majority of the values are obtained in static conditions, indeed, as reviewed by Eshragi et al. [91], just a few works have been dedicated to dynamic pistoning (5 out of 18 studies from 1975 to 2011). Regarding transfemoral amputees, there are only two works investigating pistoning.
2.5 The stump-socket interface: physical parameters
2.5.1 Temperature
Thermal sensation of human body is determined by external transfer mechanisms (convection, radiation, evaporation and conduction) and by the internal thermoregulatory system. According to these mechanisms the body responds to thermal inputs by sweating or other physiological thermoregulations like vasoconstriction and vasodilatation [96], [97]. The age, sex, activity level and circadian rhythm of a person influence such response, hence thermal comfort is subjective [65], [98]. However average values of skin temperature are around 31°, which can be used as mean value for obtaining a good thermal comfort [99], [100]. After amputation, the thermoregulatory system tends to increase the sweating rate, since there is a reduction of skin surface, and it is covered by the socket. An increase of just 1°C÷2°C is enough to induce discomfort in amputees [101]. Temperature is one of major problems of prosthetic users. It has been reported that more than 53% of them feels discomfort due to excessive heat or sweating [58], [59], [102].
In order to obtain an optimal fit and suspension of the prosthetic device, total contact sockets with liners are the standard of care. However, the total contact and the use of inorganic materials inhibit thermal transfer mechanisms, determining an increase of skin temperature and moisture [35], [73], [93], [103]. Current prosthetic materials are characterized by poor thermal conductivity (lower than 1 W/m*K) and low moisture permeability, worsening the heat transfer [104], [105]. Moreover, pressure and shear stresses on the residual limb cause unusual local temperature fluctuations in the skin and underlying tissues, by affecting the blood flow [62], [63], [67]. Consequently skin maceration and bacterial invasion may occur, producing skin irritation, friction blisters and other skin problems [65], [106]. Amputees with diabetes or vascular diseases (majority of prosthetic users) have often thermoregulatory sweating abnormalities [25], [101], [107] and a reduced convective mechanism by the circulatory system in the residual limb [45].
Perry et al. [108] analyzed skin temperature fluctuations within the socket of 5 transtibial amputees during different activities and at 14 different locations, by using thermistors. The temperature had a mean initial value of 31.4° and increased to 0.8° by staying in a seated position for 15 minutes, and to 1.7° by walking for 10 minutes. Huff et al. [106] did a new analysis on a transtibial amputee and found that the skin temperature decreased of 0.4° from its initial value, by resting for one hour, whereas incremented of 3.1° after walking for one hour. Both the studies
revealed the skin temperature lower at areas with low perfusion, such as the anterior proximal locations, and warmest at areas with high perfusion, as across the posterior section. Near large muscles metabolism and perfusion are greater than near bony regions. There are different solutions able to minimize the overheating and moisture, however there is still no satisfactory commercial systems.
Solutions to optimize stump overheating and moisture
Overheating and sweat can be avoid adopting different strategies, involving the materials, the design of the socket, and cooling devices. There are also some treatment options against perspiration including anti-perspirant sprays, Botulinum toxin, medication, electrical stimulation and surgery. However, these treatments are invasive and often not successful. In addition they may cause further side effects.
Hence, some studies have been focused on the properties of breathable fabric materials, developing liners with bacteriostatic fibers and silver ions. These avoid the bacterial invasion and unpleasant odors, and their good capillary effects eliminate the perspiration [105]. Nanomaterials are being studied as well in order to be embedded in liners and increase their thermal conductivity, but their safety at the skin interface has not been evaluated [101], [109]. Cooling vests can be also obtained by adding phase change materials, which can absorb or release heat according to their change phases [110]. However, these systems are very hard to control. Otherwise layers with highly conductive fibers can be used in order to redistribute heat and let concentrate air while another canalized layer allows perspiration escape [111].
Sockets with fenestrations are widely used in upper limb prosthetics, since there is a less strict need of liners. These sockets allow better ventilation and moisture escape [112]–[114]. Frame designs can be applied to lower limb prostheses too, but breathable liners are needed in order to exploit the advantages of this kind of solution [115] (Figure 16).
Figure 16: Solutions comprising breathable liners, modified sockets and liners. (a) Patent of a fenestrated socket and breathable liner [115], (b) perforated commercial liner for moisture escapement by a vaccum pump inside the EchelonVAC foot.
The main drawback of fenestrated sockets is the risk to enhance stresses at the brims.
Otherwise, the shape changes can be applied to the liner. The use of liners with drilled perforations along the length and at the distal end has been suggested [116]. The system let air escaping and moisture elimination by an external valve (e.g. Silcare Breathe Liner of Blatchford Inc) (Figure 16).
Many research groups focused on the development of thermoregulatory devices to be embedded in the socket or liner. The simplest system incorporates one or more thermoelectric coolers (TEC) [117], but it has low performances. Ghoseiri et al. [100], [118] developed a device that gets data from temperature sensors, sends them to a microcontroller board and compares them with a set value. The comparison is then used to control the cooling or heating function of a thermal pump. Han et al. [99], [119] proposed the use of heat pipes embedded into the socket wall and facing the liner. The first system concentrates the heat flux from the liner to a heat sink, where a compact fan moves the heat flux from the heat sink to the ambient. The second work proposes an embedded flow channel array to transfer the heat from the pipes to a cooler phase change material (ice pack) (Figure 17).
Figure 17: (a) and (b), devices developed by Web and Davis and exploiting the flow of air and fluid respectively, through channels embedded in the socket walls, (c) device developed by Han et al to dissipate
the heat with flow channales and phase change materials.
Varying the fan power and the velocity of flow [120] in the two works respectively, it was possible to compensate different activity levels. Following the same idea, Zhe et al. [121] proposed an integrated cooling system which comprises a heat pipe to focus the heat flux through a working fluid to a heat sink, regulated by a control system and heat sensors. The evaporation of the fluid draws heat and subsequently the heat sink decreases its temperature. The device includes a vacuum system in the heat pipe for changing the boiling point of the fluid with respect to that at atmospheric conditions, and a fan for blowing air across the heat pipe. On the other hand, Webber and Davis [122] suggested a socket incorporating an helical cooling channel, where air was let flowing instead of liquids. However, these cooling systems based on convective capabilities requires high power supplies for circulation, making hard to integrate them into a socket [119].
2.5.2 Stresses
As previously stated, loads at the stump-socket interface are a delicate matter, since they can lead to discomfort and skin problems [59]. Hence, possible threshold values can help avoiding these consequences [123]. Generally, even if the discomfort felt depends on the amputee’s weight, residuum tonicity and life style [124], high stresses are less tolerated at the bony prominences, because of the absence of cushioned protection by the soft tissues [58], [125], [126].
Current sensing systems are able to monitor mainly pressures [127]. Some research prototypes have been developed to evaluate shear stresses too, but they are not reliable and there are not commercial solutions for socket measurements. Regarding pressures, literature reports values of 40 kPa, as pressure threshold caused by sitting in healthy people [50], or 50 kPa and 120 kPa for
fossa popliteal and patellar tendon respectively [128]. However, common threshold values within a socket are not yet known, since they depend on the underlying tissues and on the features of the prosthesis itself.
Transtibial pressure values
Besides finding pain thresholds around the residual limb, the knowledge of the stress distribution at the socket-stump interface can help comparing different designs and suspension systems. Normal and shear stresses have been reviewed in transtibial sockets by Safari et al. [57] and Al-Fakin et al. [127]. These studies showed that in PTB sockets the average peak pressure ranges from 25 kPa to 320 kPa, with peak pressures of 415 kPa. The average peak shear stress ranges from 5 kPa to 61 kPa at the medial tibia. However, these values vary for different liners and their thickness. Boutwell et al. confirmed a reduction of pressures at the fibula head of 26 % ± 21% and an increase of ground reaction force loading peak of 3 ± 3% for thick liners [129]. Pressures change
also on the base of the suspension. The TSB socket with sleeve suspension showed greater pressures during the stance phase compared to VAS, but lower during swing phase [127]. Ali et al. [66] compared TSB socket with a pin-lock liner and a Seal-In liner, during stair ascent and descent. Results showed that the pin-lock system produces average peak pressures at the anterior, posterior and medial areas lower than the Seal-In system, as confirmed by Eshraghi et al. [130]. Eshraghi et al. compared the pressure distribution between Seal-In, pin-lock and a magnetic suspension system during locomotion and slopes. The Seal-In liner showed the greatest pressure during ascending/ descending, while the new magnetic suspension revealed the lowest pressures, reporting values ranging from 50 to 60 kPa (Table 2).
TSB + Pin-lock liner
[131], [132]
Anterior Posterior Lateral Medial
89.89 Peak 47.22 Peak 31.65 Peak 39.21 Peak
Proximal Distal Proximal Distal Proximal Distal Proximal Distal
56.10 peak upstairs 59.11 peak downstairs 57.42 peak up slope 52.22 peak down slope 58.03 peak upstairs 54.11 peak downstairs 50.15 peak up slope 60.81 peak down slope 57.10 peak upstairs 52.10 peak downstairs 59.64 peak up slope 57.34 peak down slope 54.01 peak upstairs 58.16 peak downstairs 51.73 peak up slope 53.80 peak down slope 58.31 peak upstairs 60.42 peak downstairs 56.86 peak up slope 49.32 peak down slope 63.13 peak upstairs 55.15 peak downstairs 62.99 peak up slope 56.34 peak down slope 45.56 peak upstairs 45.05 peak downstairs 44.16 peak up slope 44.26 peak down slope 43.03 peak upstairs 43.35 peak downstairs 39.14 peak up slope 39.82 peak down slope TSB + magnetic-lock liner [131], [132]
79.26 Peak 26.01 Peak 38.07 Peak 27.41 Peak
Proximal Distal Proximal Distal Proximal Distal Proximal Distal
48.30 peak upstairs 46.01 peak downstairs 48.21 peak up slope 45.21 peak down slope 50.15 peak upstairs 41.70 peak downstairs 45.02 peak up slope 53.50 peak down slope 47.42 peak upstairs 42.23 peak downstairs 49.54 peak up slope 48.82 peak down slope 43.20 peak upstairs 49.06 peak downstairs 43.71 peak up slope 55.04 peak down slope 42.78 peak upstairs 56.70 peak downstairs 54.18 peak up slope 51.13 peak down slope 45.05 peak upstairs 50.17 peak downstairs 61.32 peak up slope 54.41 peak down slope 43.15 peak upstairs 41.70 peak downstairs 42.31 peak up slope 41.02 peak down slope 41.21 peak upstairs 38.91 peak downstairs 40.08 peak up slope 40.61 peak down slope TSB + Seal-In liner [131], [132]
119.43 Peak 65.29 Peak 53.50 Peak 52.55 Peak
Proximal Distal Proximal Distal Proximal Distal Proximal Distal
69.02 peak upstairs 65.61 peak downstairs 71.14 peak up slope 67.22 peak down slope 64.04 peak upstairs 67.05 peak downstairs 63.67 peak up slope 74.20 peak down slope 80.40 peak upstairs 82.14 peak downstairs 81.66 peak up slope 72.07 peak down slope 59.10 peak upstairs 68.56 peak downstairs 65.28 peak up slope 83.00 peak down slope 61.13 peak upstairs 55.45 peak downstairs 66.89 peak up slope 55.67 peak down slope 60.01 peak upstairs 57.30 peak downstairs 69.56 peak up slope 61.19 peak down slope 52.25 peak upstairs 54.20 peak downstairs 63.95 peak up slope 49.63 peak down slope 52.20 peak upstairs 50.24 peak downstairs 60.83 peak up slope 54.11 peak down slope TSB + 3 mm thick liner [133]
Proximal Anterior Proximal Lateral Distal Anterior
237 peak mean stance 352 peak mean stance 278 peak mean stance TSB + 9 mm
thick liner [133]
177 peak mean stance 254 peak mean stance 254 peak mean stance
TSB + 6mm thick liner
[134]
Proximal Anterior Proximal Posterior Med
Lateral Med Medial Distal Anterior
166.2 Peak 215.8 peak upstairs 189.8 peak downstairs 172 up slope 165.7 down slope 184.1 non-flat road 182.7 Peak 165.5 peak upstairs 186.1 peak downstairs 170.4 up slope 190.6 down slope 184.7 non-flat road 141.7 Peak 123.9 peak upstairs 139 peak downstairs 128.1 up slope 129.6 down slope 133.5 non-flat road 76.6 Peak 68.9 peak upstairs 81.2 peak downstairs 65.6 up slope 71.5 down slope 75.7 non-flat road 175.9 Peak 140.3 peak upstairs 175.3 peak downstairs 124.5 up slope 170.9 down slope 155.4 non-flat road TSB + VAS [52]
Proximal Mid Medial Distal
Medial Mid Lateral Distal Lateral
81.5 peak stance 76.7 peak stance 67.6 peak
stance 89.5 peak stance 84.8 peak stance −26.5 average swing; −36.3 peak swing
TSB with compliant areas + VAS
[135]
Distal Proximal Lateral
88.4 peak mid stance 33.9 peak terminal stance
Anterior Proximal Medial Proximal Posterior Distal Lateral Proximal
PTB [136], [137]
41.46 average early stance 46.95 average mid stance 63.37 average late stance
59.54 average early stance 67.45 average mid stance 70.87 average late stance
56.91 average early stance 62.62 average mid stance 78.83 average late stance
53.91 average early stance 54.63 average mid stance 57.83 average late stance
