The peak principal compression strain in the muscle flapof our TTA patient was slightly higher than physiologicalcompression levels previously reported for the gluteusmuscle of healthy subjects during sitting (Linder-Ganzet al., 2007). However, the peak principal tensile strain andshear strain found in her flap were 5-fold the physiologicallevels in Linder-Ganz et al. (2007). It has been widelyreported that the residual limb is subjected to high externalshearing forces at the limbsocket interface (Zhang et al.,1998, Sanders and Daly, 1993; Quesada and Skinner,1991). It is also known that with increasing shear, thepressure necessary to occlude skin blood flow drops as lowas half the baseline value (Bennett et al., 1979; Linder-Ganzand Gefen, 2007). Hence, although compression strainsfound in the muscle flap of our subject appeared to be atphysiological levels, the high tensile and shear strains mightbe a risk for tissue viability.
Many materials are suitable for socket fabrication. As is the case with other levels of lower-limb amputation, the most commonly utilized socket material is a rigid thermosetting resin: polyester or acrylic. An increasing trend toward more flexible thermoplastic materials is evident, as in other aspects of prosthetic practice. One of the authors (J.W.M.) has fitted more than two dozen polypropylene/polyethylene copolymer sockets for hip-level amputation over the past decade with good long-term results in durability, comfort, and patient acceptance (Fig 21B-14.).
Although the anatomic differences between hip disarticulation and transpelvic (hemipelvectomy) amputations are considerable, prosthetic component selection and alignment for both levels are quite similar. The major differences are in socket design and will therefore be discussed in some detail. A full surgical report identifying muscle reattachments along with postoperative radiographs can be extremely valuable during the initial examination of the amputation site, particularly if any portions of the pelvis have been excised. This information, combined with a thorough physical examination and a precise plaster impression, will influence the ultimate fit and function of the prosthesis.
Many authors have noted that the rejection rates for lower-limb prostheses are the highest at these proximal levels. The energy requirements to use such prostheses has been reported to be as much as 200% of normal ambulation. At the same time, the lack of muscle power at the hip, knee, and ankle/foot results in a fixed, slow cadence. As a practical matter, only those who develop sufficient balance to ambulate with a single cane (or without any external aids at all) are likely to wear such a prosthesis long-term. Those who remain dependent on dual canes or crutches for balance eventually realize that mobility with crutches and the remaining leg, without a prosthesis, is much faster and requires no more energy expenditure than using a prosthesis does.
He specializes in the
diagnosis and treatment of:
Congenital lower limb deformities (eg., congenital femoral deficiency [PFFD], hemihypertrophy, fibular hemimelia, tibial hemimelia,
posteromedial tibial bow, congenital dislocation of patella, congenital pterygium of knee, congenital pseudarthrosis)
Congenital upper limb deformities (e.g., short humerus and forearm, radial clubhand, ulnar clubhand, radiohumeral synostosis,
radioulnar synostosis, syndactyly, absent thumb, polydactyly, congenital pterigium of elbow)
Other upper and lower limb deformities due to: growth arrest, fractures, radiation, infection.
Post-traumatic limb deformities and leg length discrepancies (e.g., malunion)
Bone healing problems (e.g., problem fractures, delayed union, nonunion, congenital pseudoarthrosis)
Bone defects, bone and joint infections (e.g., osteomyelitis, septic arthritis, sequellae of neonatal sepsis)
Skeletal dysplasias (e.g., achondroplasia, hypochondroplasia, spondyloepiphyseal dysplasia, pseudoachondroplasia,
chondrometaphyseal dypslasia, mesomelic dysplasia, Ellis-van Creveld, melorheostosis, diastrophic dwarfism)
Tumor-like conditions (e.g., fibrous dysplasia, neurofibromatosis, multiple hereditary exostoses (MHE), Ollier’s disease)
Metabolic disorders (e.g., rickets, Paget’s disease, osteogenesis imperfecta, osteopetrosis)
Other miscellaneous developmental deformities (e.g., Blount’s disease, growth arrest, neonatal sepsis sequelae)
Joint contractures and joint stiffness (e.g., Perthes disease, knee and ankle arthritis)
Foot deformities (e.g., clubfoot, vertical talus, dropfoot, equinus, flatfoot, short metatarsals (brachymetatarsia), Charcot-Marie-
Hip dysplasia in young adults and neonatal septic hip sequellae
Short residual limb following amputation
Joint preservation for arthritis of the hip, knee and ankle
Peripheral nerve disorders (e.g., nerve entrapment)
Constitutional short stature
He is currently completing a book on the Congenital Lower Limb Deformities.
He was professor of Orthopedics, chief of Pediatric Orthopedics and co-director and founder of the Maryland Center for Limb
Lengthening and Reconstruction at the University of Maryland between 1987-2001.
The friction-brake stance control (safety) knee is probably the second most frequently utilized component. Because there is very little increase in cost or weight and reliability has been good, many clinicians feel that the enhanced knee stability justifies this approach, particularly for the novice amputee. Missteps causing up to 15 degrees of knee flexion will not result in knee buckle, which makes gait training less difficult for the patient and therapist. The major drawback to this knee is that the limb must be non-weight bearing for knee flexion to occur. Although this generally presents no problem during swing phase, some patients have difficulty in mastering the weight shift necessary for sitting. It should be noted that use of such knee mechanisms bilaterally must be avoided. Since it is impossible for the amputee to simultaneously unload both artificial limbs, sitting with two stance control knees becomes nearly impossible.
A third type that has proved advantageous for this level of amputation is the polycentric (four-bar) knee. Although slightly heavier than the previous two types, this component offers maximum stance-phase stability. Because the stability is inherent in the multilinkage design, it does not erode as the knee mechanism wears during use. In addition, all polycentric mechanisms tend to "shorten" during swing phase, thus adding slightly to the toe clearance at that time. Many of the endoskeletal designs feature a readily adjustable knee extension stop. This permits significant changes to the biomechanical stability of the prosthesis, even in the definitive limb. Because of the powerful stability, good durability, and realignment capabilities of the endoskeletal polycentric mechanisms, they are particularly well suited for the bilateral amputee. Patients with all levels of amputation, up to and including translumbar (hemicorporectomy), have successfully ambulated with these components.
Objective. The aim of this study is to evaluate residual muscle function abnormalities after total knee replacement, with respect to gait kinematics and kinetics. Design. Longitudinal study on a follow-up of up to two years. Background. Gait usually presents an excellent improvement after total knee replacement. Nevertheless, some kinematics and kinetics abnormalities persist even after a long period of time and they might have implications in long-term prosthesis failure. Additionally, lower limb muscle activity has not been sufficiently studied in the past directly by means of dynamic EMG. Methods. Nine patients who had a posterior cruciate sparing total knee replacement design were evaluated by means of clinical assessment and gait analysis at the end of rehabilitation trials at six, twelve and twenty four months. EMG from trunk and lower limb muscles was registered and elaborated through a statistical detector for the on-off timing. Results. Gait analysis showed a slow gait, with a "stiff knee gait pattern" and prolonged muscular co-contractions during stance. Conclusions. Knee kinematics and kinetics abnormalities during loading acceptance after total knee replacement are associated with co-contractions in muscular activation pattern.
Most trans-tibial amputation (TTA) patients use a prosthesis to retain upright mobility capabilities. Unfortunately, interaction
between the residual limb and the prosthetic socket causes elevated internal strains and stresses in the muscle and fat tissues in the
residual limb, which may lead to deep tissue injury (DTI) and other complications. Presently, there is paucity of information on the
mechanical conditions in the TTA residual limb during load-bearing. Accordingly, our aim was to characterize the mechanical conditions
in the muscle flap of the residual limb of a TTA patient after donning the prosthetic socket and during load-bearing. Knowledge of
internal mechanical conditions in the muscle flap can be used to identify the risk for DTI and improve the fitting of the prosthesis. We
used a patient-specific modelling approach which involved an MRI scan, interface pressure measurements between the residual limb and
the socket of the prosthesis and three-dimensional non-linear large-deformation finite-element (FE) modelling to quantify internal soft
tissue strains and stresses in a female TTA patient during static load-bearing. Movement of the truncated tibia and fibula during loadbearing
was measured by means of MRI and used as displacement boundary conditions for the FE model. Subsequently, we calculated
the internal strains, strain energy density (SED) and stresses in the muscle flap under the truncated bones. Internal strains under the tibia
peaked at 85%, 129% and 106% for compression, tension and shear strains, respectively. Internal strains under the fibula peaked at
substantially lower values, that is, 19%, 22% and 19% for compression, tension and shear strains, respectively. Strain energy density
peaked at the tibial end (104 kJ/m3). The von Mises stresses peaked at 215 kPa around the distal end of the tibia. Stresses under the fibula
were at least one order of magnitude lower than the stresses under the tibia. We surmise that our present patient-specific modelling
method is an important tool in understanding the etiology of DTI in the residual limbs of TTA patients.
r 2008 Elsevier Ltd. All rights reserved.
Prosthesis; Deep tissue injury; Pressure ulcer; Patient-specific finite element model; Rehabilitation