Approximately 50% of the elderly population, and even young patients normally after traumatic events, suffer joint disorders. In fact, the World Health Organization (WHO) declared the years 2000-2010 to be the bone joint decade. Cartilage damage, osteochondral defects (defects going from cartilage through bone) and cartilage degeneration is thus an important topic requiring deeper understanding. Mechanical conditions influencing osteochondral healing is the theme of the present work. The mechanical quality of the repaired tissues after the healing process is essential to preserve joint functionality. Tissue repair means restoration of the initial mechanical quality of affected tissues (after injury or trauma) and the consideration of biological aspects related with their healing process. Successful healing is characterized not only by reestablishing full functionality but also by avoiding long-term acceleration of degenerative damage. The tissues that actually are most frequently treated for repair are skin (replacement), cartilage (induced repair, replacement), bone (replacement), liver (artificial transplants) and heart (cardiac prostheses) (Lysaght and Hazlehurst, 2004). To be able to repair cartilage tissue, the study of its biomechanical aspects is required.
Biomechanics can be defined as the application of principles derived from engineering sciences (mechanics) to understand and to explain biological events (gait analysis, musculoskeletal movement) and processes (fracture healing, osteochondral healing, damage, cell differentiation). Classical concepts of physics, chemistry, mathematics and biology are also used. From this topic, tissue engineering was developed as a research discipline in the 1970s and matured into preclinical application during 1990s (Lysaght and Reyes, 2001). It has been defined as devices or processes that: 1. Combine living cells and biomaterials. 2. Utilize living cells as therapeutic or diagnostic reagents 3. Generate tissues or organs in vitro for subsequent implantation and/or 4. Provide materials or technology to enable such approaches (Griffith and Naughton, 2002;Lysaght and Hazlehurst, 2004;Lysaght and Reyes, 2001).
A vigorous development of tissue engineering has been focused on repairing osteochondral defects. The mechanobiological study of osteochondral defect healing requires the knowledge of physiology and pathology of cartilage. Articular cartilage acts mainly by transmitting physiological loads between joints to the underlying subchondral bone. In an intact situation, articular cartilage shows a white, smooth and almost frictionless surface. When damage occurs, the first changes appear in the surface: micro-cracking, fibrillation, loss of the superficial layer, osteophytes formation, which implies an increase of the friction coefficient and diminution of the range of motion. Moreover, a reduction of the water content can be observed. Finally, the tissue could fail, thereby creating osteochondral defects (Buckwalter and Mankin, 1998;Buckwalter and Mankin, 1998;Buckwalter and Mankin, 1998;Carter, et al., 2004;CHEN and BROOM, 1998;Eckstein, et al., 2001). As a result, understanding biological (Benjamin and Ralphs, 2004;Revell and Heatley, 1988;Zhurakovskii, et al., 2002), genetic (Fortier, et al., 2001;Grande, et al., 2003;Hidaka, et al., 2003;Pascher, et al., 2004), mechanical (Boschetti, et al., 2004;D'Lima, et al., 2001;Frost and Jee, 1994;Heiner and Martin, 2004;Kääb, et al., 1998;Li, et al., 2005;Smith, et al., 2004) and mechanobiological aspects (Beaupré, et al., 2000;Carter, et al., 2004;Carter and Wong, 2003;Frederick H. Silver and Bradica, 2002;Loboa, et al., 2003;Raimondi, et al., 2001;Sarin and Carter, 2000;Silver and Bradica, 2002;van der Meulen and Huiskes, 2002) of cartilage and osteochondral healing becomes important. Specifically, research has been directed to find a practicable clinical solution to restore joint mechanics in order to avoid or to delay the employment of prostheses for total joint replacement. Interdisciplinary cooperation is necessary to develop tools that can be employed to diagnose joint damage, to repair it and to maintain the optimal long-term mechanical properties of the repaired tissues (Buckwalter and Brown, 2004;Buckwalter, et al., 2003;Butler, et al., 2004;Guilak, et al., 2001;Hattori, et al., 2004;Poole, 2003;van den Berg, et al., 2001;Wang and Yu, 2004). Actual studies (animal models, numerical approaches, etc.) investigate the role of mechanical conditions on osteochondral healing. The loads acting on the joint should then be evaluated as well as the mechanics of the bone to facilitate a more precise local study of the joint region.
The aims of this project were to understand the biomechanical behavior of an intact and injured joint under physiological loads while taking into account the influence of the bone quality, and subsequently to study the mechanical aspects of osteochondral healing.
This work is subdivided into five chapters. In Chapter 1, the introduction, four sections are found: The clinical problem, related to osteochondral healing is described. Cartilage analysis and function follows in which the physiology of the articular cartilage, the pathology of synovial joints and its treatments, as well as the techniques used to measure the cartilage material properties are presented. Additionally, this section describes several tissue differentiation models. The third section, preliminary work summarizes the previous two projects by Lill, related to bone – joint mechanics and Bail, related to osteochondral healing, whose results were used to validate the models developed in the present work. Hypotheses and aims of the present project are then stated. Subsequently, the remaining chapters, material and methods, results, discussion and conclusions are presented.
Trauma, degenerative diseases or repetitive and sudden loads, through sports activities for example, can produce articular damage, which could result in osteochondral defects (Arokoski, et al., 2000;Brittberg and Winalski, 2003;Petersen, et al., 2003;Sgaglione, 2003). Approximately 60% of these defects occur after trauma in teenagers (Buckwalter, 1999) (Fig. 1.1). In older patients, articular damage is attributable to degenerative diseases (50% in individuals > 60 years old) (Buckwalter and Mankin, 1998;Hjelle, et al., 2002).
Independently of the causes (trauma or disease), osteochondral defects commonly do not heal or are filled with a fibrous instead of hyaline cartilage, which is characterized by reduced mechanical properties compared with the highly specialized compressive capabilities of the native tissue. As a result, fragmentation and degradation of the fibrous tissue may occur, thus exposing the subchondral bone to further damage (Buckwalter, 1999). Without treatment, osteoarthrosis and potentially joint replacement could be expected (Petersen, et al., 2003). The treatment of chondral and osteochondral defects remains one of the most ambitious challenges in the field of traumatology and orthopedics.
|Fig. 1.1: Schematic representation of the relationship between age and types of chondral defects. Osteochondral defects included osteochondritis dissecans and low energy osteochondral fractures occur most frequently in skeletally immature individuals and individuals who recently have completed skeletal growth. Chondral defects, localized tears or disruptions of articular cartilage that do not extend into subchondral bone occur most frequently in young adults. With increasing age the frequency of degenerative lesions of the articular surface increases (Buckwalter, 1999) (figure with corresponding text).|
Some experimental studies with animals have demonstrated the influence of mechanical conditions on osteochondral repair (Buckwalter, et al., 2003;Case, et al., 2003;Salter, et al., 1980). In an in vivo experiment with 120 rabbits, 480 osteochondral defects were surgically created and stimulated to induce repair through passive motions (Salter, et al., 1980). The results demonstrated that healing was not only faster than in the control group (without stimulation) but also that after histological analysis the restored tissues showed a better mechanical quality.
More recently, tissue differentiation models have been developed to analyze the influence of specific mechanical boundary conditions on healing. The usage of these models allows the determination of the strain environment for differentiation that enables healing, which is indeterminable in vivo (Duda, et al., 2005;Kelly and Prendergast, 2004;Kelly and Prendergast, 2004). Improved understanding of the influence of mechanical conditions on osteochondral repair may contribute to elucidate its relation with the biological healing process in order to prevent osteoarthrosis and thus, limit loss of joint functionality (Buckwalter and Brown, 2004). Apparently, clinical reports indicate that convex joints are more frequently affected by osteochondral defects compared to concave ones (Hjelle, et al., 2002). However, no study has so far been published to explain this occurrence. Healing of osteochondral defects could be influenced by the local mechanical environment generated after changes in the mechanical boundary conditions of the defect.
Considering the restricted capacity of the cartilage to repair itself, techniques have been developed to reestablish joint functionality (Cancedda, et al., 2003;Chu, et al., 1995;Wang, et al., 2004). One of these treatments employs the usage of autografts. Since only a limited quantity of unloaded cartilage regions are available, the implantation of predesigned plugs made out of biomaterials becomes an important alternative (Buckwalter, 2003;Evans, et al., 2004;Giannini, et al., 2002;Gole, et al., 2004). The mechanical properties of the used fillings and the subchondral bone possibly need to be similar in order to reduce stress concentration at the interface graft-host tissue. This could avoid the frequently reported loss of anchorage between the host tissue and the defect filling (Akens, et al., 2001;Hangody and Fules, 2003;Kuroki, et al., 2004;Nam, et al., 2004;Tibesku, et al., 2004).
From a study of 1000 arthroscopies, 61% of the revised patients revealed any type of chondral or osteochondral defects of which 19% were focal osteochondral defects (Hjelle, et al., 2002). In the USA, the annual costs for the treatment of joint degeneration amounts to more than $60 billion, thus affecting the quality of life of more than 20 million of Americans, implying an immense social and economic burden1 (Buckwalter, et al., 2004).
As a result, a high percentage of clinical and basic research has aimed to analyze the nature of the cartilage and the surrounding tissues of the joint. After interdisciplinary studies of the joint region, it has been possible to elucidate its mechanical behavior, its structure characterization, its material properties and its diseases. The knowledge of cartilage nature and properties can be used to prevent cartilage damage or to promote its repair. Even at a cellular level, the mechanical properties of chondrocytes have been analyzed. The most frequent techniques used are high-resolution sonography (Keogh, et al., 2004), atomic force microscopy (AFM) (Park, et al., ), nuclear magnetic resonance (NMR) (Perez and Santos, 2004), magnetic resonance imaging (MRI) (Burstein and Gray, 2003;Chen CT, et al., 2003;Graichen, et al., 2004;Kornaat, et al., 2004) and Doppler effect (Strunk, et al., 2004). Some of these techniques have been implemented in innovative devices (ultrasound arthroscopic indenters (Laasanen, et al., 2003;Toyras, et al., 2001), creep cytoindentation apparatus (CCA) (Koay, et al., 2003), handheld dynamic indenters (Appleyard, et al., 2001), arthroscopic dynamic and static indenters (Lyyra-Laitinen, et al., 1999;Toyras, et al., 2001)) to measure in vivo and in vitro mechanical parameters (elastic modulus of Young, permeability, Poisson’s ratio, etc.) from cartilage specimens or chondrocytes. These mechanical parameters show how cartilage changes its material properties as a response under specific physiological and pathological conditions. Additionally, new concepts such as pathways signaling and regulating chondrocyte activity have been used to explain the complex mechanobiological behavior of this zone. In this work a model for tissue differentiation is proposed as a potential tool to understand the mechanical aspects of this region. Using the finite element method (FEM), changes in the mechanical properties of the tissues during healing could be predicted in dependence of the acting mechanical conditions. For the development of a tissue differentiation model, some concepts involving bone remodeling, tissue differentiation and cartilage repair were linked and the corresponding models studied.
A synovial joint consists of hyaline cartilage, subchondral bone plate (calcified cartilage), subchondral bone and cancellous bone. Hyaline cartilage is an avascular, aneural and alympathic tissue, which lives from diffusion. The synovium, a fine membrane at the articular interfaces, secretes synovial fluid and provides the nutrients required by the tissues within the joint. Articular cartilage is a multiphasic material with a fluid and a solid phase. The fluid phase is composed of water and electrolytes and the solid phase consists of chondrocytes, collagen, proteoglycans and other proteins. Each phase contributes to the mechanical and physicochemical properties of the articular cartilage. Physical models should thus describe cartilage as a biphasic material, whose mechanical behavior is given by the concepts and laws used in soil mechanics (rule of mixtures). Mow and his group have made several contributions to illustrate this and justify the usage of this assumption (Mow, et al., 1989;Mow and Ratcliffe, 1997).
Structurally, cartilage components (chondrocytes and extracellular matrix) are arranged in distinct zones each with a different and specific structure and function: a superficial or tangential zone, a middle zone or transitional zone, a deep zone and a zone of calcified cartilage (Fig. 1.2) (Buckwalter, 1983;Horký, 1993;Hunziker, et al., 2002). These zones are characterized by a distinct microstructure associated with a specific capacity to support and to transmit acting external loads. Hence a relation between each structure zone of the cartilage and its function has been established. The superficial tangential zone (10 to 20% of the total cartilage thickness) is limited by a superficial layer or lamina splendes, which serves as a skin or barrier protecting the cartilage from the synovial environment. This zone has a high water content and shows ellipsoidal chondrocytes parallel to the synovial surface. This structure contributes principally to the dissipation and transmission of load, principally tensile and shear, allowing large compressive strains. The middle zone represents approximately 40% to 60% of the total cartilage thickness with low fluid flow and shows redounded chondrocytes forming a quadratic array similar to vertical “arcades” (Benninghoff model). Such structural arrengements are appropriate for the support and transmission of compressive loads allowing moderate compressive strains (Mow and Ratcliffe, 1997).
|Fig. 1.2: Schematic representation of the cartilage zones. The mechanical properties of cartilage are changing in a functional – structure relation. The elastic modulus of Young is reduced from the upper zone (tangential) to the deeper regions (Mow and Ratcliffe, 1997).|
The deep zone (about 30% of the total articular thickness) shows a minor quantity of chondrocytes, which are aligned, perpendicular to the subchondral bone traversing the “tidemark” or lamina limitants with anchorage in the calcified cartilage. The structural conformation of all these zones has been associated with the cartilage’s capacity to support compressive loads. However, according to latest publications minor differences have been encountered, when these zones have been included or excluded in the simulation of the mechanical behavior of the joint. In this project, cartilage was simulated as biphasic, isotropic, linear elastic without consideration of the above described structural cartilage zones (tangential, middle, and deep zone). Only insufficient information exists to describe all mechanical parameters of each cartilage zone necessary to model it as a non-homogeneous biphasic material (e.g. permeability and porosity defined in dependence of the cartilage deep). Therefore, cartilage was modeled as an isotropic biphasic material - that is, stiffness, Poisson’s ratio and permeability were defined to be the same in each material direction. However, the differentiation model created in this project is able to predict inhomogeneous cartilage stiffness. That is, the stiffness distribution of the cartilage after healing was different in different locations of the cartilage (non-homogeneous). In this form, although the tissue differentiation model started with a homogeneous material definition for the cartilage, a non-homogeneous stiffness distribution can be predicted.
When an external load acts on a joint, the fluids inside the cartilage are activated and its movement known as fluid exudation is initiated. Fluid exudation implies the movement of the fluids in a parallel but opposite direction to the applied load. The principal tasks of these fluids are to attenuate the acting loads between the collagen matrix and the subchondral bone (Bader, et al., 1992;Broom and Poole, 1982;Buckwalter, 1983;Buckwalter and Mankin, 1998;Ghivizzani, et al., 2000;Hunziker, et al., 2002;Hunziker, et al., 1996;Ikenoue, et al., 2003;Poole, et al., 2001). The collagen fibrils query the form and tensile strength properties to the cartilage tissue. The interaction between the aggrecans (a component of the cartilage) with the cartilage water is responsible for its compressive stiffness and its durability. When the loads are slowly applied, the cartilage is able to achieve a maximal deformation by fluid exudation and fluid redistribution, which reduces the effect of the acting load by generation of hydrostatic pressures. In contrast, when sudden loads act on the joint, a high percentage of the acting load is directly transmitted to the collagen matrix and could cause its collapse (fractures). Finally, cartilage cells are affected: overstrains could cause structural changes producing cellular damage including cell death. Consequently, joint mechanics are altered, thus loosening their ability to attenuate and to transmit the acting loads. In normal conditions sudden and quick external forces are absorbed by muscle contractions as well. However, in some situations the loads act so rapidly that neither are the muscles able to respond nor are the fluids inside the cartilage able to attenuate the acting forces. Damage at the articular surface occurs. A study of the response of human’s articular cartilage from humans to blunt trauma showed that it could resist impact loads of as much as 25N/mm2 (25MPa) without apparent damage. However, repetitive loading can propagate along vertical cartilage fissures from the joint surface to the calcified cartilage, extending oblique fissures into areas of intact cartilage. Thereby the affected areas increase, eventually generating cartilage flaps and free fragments (osteochondrosis dissecans). In general, three major cases exist in which damage of cartilage and joint degeneration may occur: trauma, disease or sports.
Three classes of chondral and osteochondral injuries can be identified based on the type of tissue damage and repair response: 1. Damage of the joint surface without visible mechanical disruption of the articular surface; however this does cause chondral damage and may cause damage of the subchondral bone. 2. Mechanical disruption of the articular surface limited to articular cartilage 3. Mechanical disruption of articular cartilage and subchondral bone simultaneously (Buckwalter, 2002). In this project, osteochondral defects corresponding to the third injury category were analyzed.
To classify articular defects, some systems and nomenclatures have been suggested: Outerbridge described four grades of cartilage damage in chondromalacia. Grade I describes softening of the surface, the grade II fissuring without reaching the subchondral bone, and grades III and IV describe defects going into or beyond the subchondral bone plate with exposition of the bare bone. Bauer and Jackson used a descriptive arthroscopic classification. Grades I to IV are used for defects on the articular cartilages with cracks, flaps and crater formation and grades V and VI for acute injuries and degrading cartilage with fissuring for older degenerative lesions (Nehrer and Minas, 2000). In order to standardize the system of articular defect classification, the International Cartilage Repair Society (ICRS) proposed a nomenclature to describe the grade of damage of the cartilage (Fig. 1.3). This classification is based on a combination of the defect depth (quantitatively) and a qualitative description of each joint region. Four different grades of the defect have been defined: the first three involve different depths of defects inside the cartilage and the last grade describes a defect which moves to the subchondral bone. In all computer models developed in this project, cartilage defects grade 4B (Outerbridge IV) were simulated. However, some cases following a detailed histomorphometric analysis of cartilage defects revealed that damage of the matrix can occur without disruption of the surface. Subsequently, rupture of the collagen fibrils, depletion of proteoglycans and increase in the hydraulic permeability leading to insufficient mechanical function of the joint is expected with consequential joint degeneration.
|Fig. 1.3: ICRS classification of osteochondral defects. In this project osteochondral defects grade 4B (severely abnormal) were modeled.|
Conservative treatments such as the use of drugs could help to avoid severe pain and to reduce the inflammatory response of the cartilage damage. Physiotherapy is used to reduce the symptoms of the cartilage damage. Joint mobility and range of motion should be thereby maintained. When cartilage suffers severe damage surgical treatment is required. Although treatment of symptomatic full thickness defects remains a challenge, the objectives are to restore joint function, providing a pain-free weight-bearing joint with full range of motion. Additionally, the rebuilt tissue should demonstrate mechanical properties that prevent any articular degeneration. In general, two different treatments exist: one that simulates or initiates a repair process and another that uses replacement or transplantation of tissues. The first category includes debridement, lavage, blot clot techniques, osteotomy, joint distraction, soft tissue grafts, cell transplantation, growth factors and artificial matrices. In essence these techniques support chondrogenic cells to reach the injury site. Cells can generally migrate from the bone marrow, from a graft of periosteum or perichondrium, from cultured periosteal or perichondreal cells or from pure chondrocytes. The second category includes graft transplantations: autografts, allografts und xenografts.
Jackson, a pioneer of arthroscopy, used the effect of joint lavage for the first time in North America. He found that during arthroscopy the usage of high flow water cleans the knee joint of debris from degenerated and inflammatory regulators in the joint. This fact produced pain relief without the necessity of any surgical intervention. Some patients showed improved joint functionality after 3 years of arthroscopic examination. The technique was combined with debridement of cartilage flaps, retrieval of loose bodies and partial synovectomy. Even better results were obtained as a result. The technique was later applied for treatment of young patients. However, after comparing the healing outcome for 14 years, only 50% of the patients showed joint degeneration. Thus this technique is unable to induce any repair and therefore cannot prevent joint degeneration. Surgical treatment of joint defects includes blot clot techniques (abrasion, drilling and microfracture), osteotomy, usage of soft tissue grafts and artificial matrices to stimulate the formation of a new articular surface.
The depth of the defect is essential for cartilage repair. When the depth of the defect reaches the subchondral bone disruption of the blood vessels occur. Blood clots inundate the defect region. If overloading is avoided, undifferentiated mesenchymal cells migrate into the clot, proliferate and differentiate to cells with morphological features of chondrocytes. The differentiation process continues until a tissue with stiffer mechanical properties fills the defect. However, it is up to now unclear exactly how this process takes place. Mechanical, biological and chemical pathways have been proposed and analyzed. Even if the initial defect does not reach to the subchondral bone, a penetration into the subchondral bone layer is artificially created in order to promote a clot blood formation and the subsequent healing cascade. These surgical techniques belong to the most frequently used in clinical practice and include subchondral drilling, microfracturing and abrasion arthroplasty.
In 1959 this technique was described first by Pridie and it is the oldest of these perforation techniques. The subchondral bone is drilled, bleeds and the healing process is started. Usually fibrous tissue is formed after the usage of this technique. Its main risks are damage to the subchondral bone by heat from the drill or causing subchondral bone haematoma. The obtained results are good, achieving an efficiency of 85% after 62 months of follow-up, principally when this technique is combined with a high tibial osteotomy.
Steadman employed this technique for the first time in 1999. After a meticulous debridement of the cartilage defect (including the calcified cartilage), the subchondral bone is penetrated using small awls approximately 3 mm deep and with perforations approximately 2 to 3 mm apart. From a study of 255 consecutive cases of subchondral defects grade IV (ICRS classification) performed by Steadman’s group, 75% of the patients showed pain relief and good functional outcome, 20% demonstrated no benefit and 5% needed revision surgery after 5 years (Steadman, et al., 2001). The advantage of this technique in comparison with drilling is that the possible damage of the subchondral bone by heat of the drill is avoided. Studies conducted by Rodrigo et al., showed that this technique, accompanied with a rehabilitation therapy of passive motion, increases its efficiency even further.
Abrasion arthroplasty is performed arthroscopically with an automated burr by removing up to 2mm of the exposed sclerotic bone down to the vasculature of the subchondral bone plate and fibrillation at the border of the defect. When the tourniquet is relaxed a blood clot is formed in the defect. This technique achieves a 53% efficiency rate in elderly patients and 86% in short-term follow-up of younger patients (Johnson, 2001). Bone marrow stimulation techniques require an intensive rehabilitation program. However, passive motion and protected weight bearing during the first six weeks have shown better healing results than cases lacking this therapy.
Although the blot clot techniques allows pain relief and an acceptable grade of joint functionality, the repaired tissue always consists of fibrous tissue only. Additionally, it was found that after repetitive intentional penetrations of the subchondral bone, the formation of multiple calluses as a repair response could cause subchondral sclerosis generating a negative cascade of biological responses that finally deteriorates the newly differentiated cartilage.
Osteotomy is a surgical technique to treat hip and knee joints with localized loss or degeneration of the articular surfaces. The surgery consists of a cut of the bone tissue in the proximal or distal region. The aim of the surgery is to perform changes in the slope or in the angle formed between the planes passing through the contact areas of the articular surfaces of the bones involved. Such changes reduce, or even eliminate, contact pressures or overloaded regions that could cause damage of the articular surfaces. These changes induce variations in the rotation centers of the hip-knee-ankle joint system correcting its possible misalignment (Buckwalter and Mankin, 1998). Alignment is measured on an axis passing through the center of the femoral to the center of the ankle (up to 1.2° varus in healthy patients). A high tibial osteotomy refers to a cut below to the level of the articular surface of the proximal tibia, normally on the tibial tuberosity, parallel to the joint, which can be complete or partial (an intact bone region remains). Patella kinematics and patello-femoral contact areas have been studied in healthy patients and in patients with genus varus and mild osteoarthritis to determine the mechanical parameter that could be altered due to improper joint alignment and how it could be corrected (Hinterwimmer, et al., 2004). Osteotomy has been used to treat primary osteoarthrosis (developed without trauma) of the ankle. In general, osteotomies can be performed in the coronal plane (varus and valgus osteotomies) or in the sagittal plane (flexion and extension osteotomies). The effect of correction in the sagittal and frontal plane by high tibial osteotomy has been demonstrated (Marti, et al., 2004;Trumble and Verheyden, 2004).
Recent studies showed that the valgus high tibial osteotomy for unicompartmental varus osteoarthritis could reduce compressive pressure distribution in the posterior part of the tibial plateau. In this form the damaged zone could be unloaded avoiding its total degeneration (Agneskirchner, et al., 2004). Other authors combine high tibial osteotomy for varus knee correction with resurfacing techniques (subchondral bone penetration) to treat chondral defects (Sterett and Steadman, 2004). In this technique, external fixators are used to stabilize the osteotomy. The question is to decide when there is “enough” consolidation of the bone for safe removal of the external fixator. This technique has the principal advantages of decreasing symptoms and stimulating the formation of new cartilage. However, the patients suffer recurrent pain and demonstrate evidence of progressive osteoarthritis.
This technique is used to treat chondral and osteochondral defects in which soft tissues (fascia, joint capsule, muscle, tendon, periosteum or perichondrium) are used to cover the defect region. Two theories underlie this clinical practice: 1. The defect should be covered mechanically and 2. In the case of periosteum or perichondrium a number of pluripotent stem cells in the perichondrium or in the cambium layer of the periosteum exist to differentiate into chondrocytes. The soft tissue grafts can be fixed into the defect using fibrin glue or sutures. Typical failures of this technique are calcification and ossification in the graft area. In this case collagen type X (bone) instead type II (cartilage) is detected indicating chondrocytes hypertrophy, which promotes enchondral ossification in the defect (Nehrer, et al., 1999;O'Driscoll, 1999).
In experimental studies with animals osteoperiosteal grafts were implanted to treat osteochondral defects. The potential of immobilization, intermittent active motion, and continuous passive motion to stimulate repair was evaluated (O'Driscoll and Salter, 1986). A group of animals without stimulation was used for control. After five weeks, hyaline cartilage was predominantly found in 70% of the defects treated with continuous passive motion compared with 10% observed in the group treated with immobilization, intermittent passive motion, and the control group.
In a later study periosteum transplants were combined with the application of transforming growth factor-ß. Experimental studies with animals showed good results. Xeno-transplantation has been studied as an alternative. In an experiment with animals, cultured chondrocytes from pigs were injected under a periosteal flap to treat chondral defects in rabbits. The chondral defect was filled to 90% with good integration between the host cartilage and the neo-differentiated cartilage. The repair tissue showed a smooth surface with cells similar to chondrocytes and a hyaline-like extracellular matrix (Ramallal, et al., 2004). However, in clinical practice detachment of the graft can occur later (O'Driscoll, 2001;O'Driscoll and Fitzsimmons, 2001;O'Driscoll, et al., 1988).
This method is complementary to the soft tissue transplantation technique. Autologous chondrocytes are cultured in vitro and are injected under the soft tissue transplant, which is sutured to the defect. The reason for the selection of chondrocytes for implantation is obvious: chondrocytes form cartilage by synthesis and chondrocytes are responsible for the unique feature of articular cartilage. The goal of the orthopedic surgeon is to try to deliver an optimal number of the acquired cell types (autologous chondrocytes) into the cartilage region to best achieve repair. Peterson and co-workers performed the first investigations using this method in 1984. Autologous chondrocytes transplantation was used to treat defects in the rabbit patella and comparisons between healing after periosteum transplantation with and without application of a suspension of cultured autologous chondrocytes were carried out. Brittberg considerably improved the surgical technique and used the same rabbit model experimentally (Brittberg, 1999;Brittberg, et al., 1994;Brittberg, et al., 2001;Brittberg and Winalski, 2003). Subsequently, the technique was employed to treat chondral defects at the knee. Brittberg and co-workers have obtained good results with patients for several years. They are convinced that the future in the treatment of chondral and osteochondral defects should include the development of biodegradable grafts with implantation of cell suspensions containing a large number of autologous chondrocytes. They called this possible technique biomedical surgery (Brittberg, et al., 2001). The O’Driscoll group has investigated this technique and compared the usage of autologous chondrocytes to human mesenchymal stem cells. They demonstrated that the usage of stem cells showed better healing outcome than by injection of autologous chondrocytes (O'Driscoll, 1999). Autologous chondrocyte transplantation is actually a promising field of research in which the question has been directed to determine the optimal type of cell (should autologous chondrocytes be replaced by autologous stem cells?) as well as the number of required cells to guarantee cartilage repair.
These techniques are frequently used to treat osteochondral defects of high diameter. They use transplantation of cartilage as a part of an osteochondral graft to replace focal regions of damaged articular cartilage. The plugs normally have a diameter of 5 to 15mm and can be used to fill defects of up to 9 cm2 of superficial area. The graft technique is subdivided in autografts, when the graft is taken from an “unloaded” region of an articular surface of the same patient, allografts when the graft is taken from another patient (same specie), and xenografts, when the graft is taken from articular surfaces of another organic entity (animals). Each technique is selected in dependence of the bone quality, age of the patient and size of the defect. The principal advantages of using grafts are to provide a fully formed articular cartilage matrix and the potential for transplanting viable chondrocytes that can maintain the matrix.
Fibrous cartilage is appropriate to support and to transmit tensile loads, but in fact, in the majority of the cases, compressive loads act on synovial joints (e.g. at the knee), which could be optimally supported by hyaline cartilage. These joints are known as weight-bearing joints. For this reason cartilage transplantations from the same patient are currently used to fill the defect using autografts taken from “unloaded” regions at the joint. Since only a limited number of these regions are available, and such a transplantation in some cases represents a risk for the patient, allografts or predesigned biomaterials with a porous structure are frequently used to fill the defect (Evans, et al., 2004;Hunziker, 1999;Jackson, et al., 2001;Porter, et al., 2004). This practice is specially employed for the treatment of osteochondral defects in young patients (Buckwalter, 1995;Roughley, 2001) or in patients with defects with a diameter smaller or equal to 20 mm (Buckwalter, 1999;Ghadially and Ghadially, 1975;Hunziker, 2002;Kaar, et al., 1998;Newman, 1998;O'Driscoll, 1998;O'Driscoll and Fitzsimmons, 2001). In critical situations the usage of transplantations of stem cells from the bone marrow or from perichondrium are recommended (Barry and Murphy, 2004;Breinan, et al., 1997;Grande, et al., 2003;Meinel, et al., 2004;Oreffo and Triffitt, 1999;Rutherford, et al., 2003;Shapiro, et al., 1993;Wakitani, et al., 1994).
The first clinical reports about the usage of autografts showed a newly formed cartilage with 70-80% of hyaline consistence. However, the principal reason to use grafts in a long-term treatment is to conserve the original structure and the mechanical quality of the subchondral bone and within its capacity to transmit the acting external forces. In the present work, the effect of predesigned biomaterials on healing was evaluated using a tissue differentiation model to simulate osteochondral repair. Other treatments include the usage of hormones, medicaments or growth factors to promote or to improve osteochondral repair.
Basically this technique involves cylinders of cartilage and subchondral bone taken from “unloaded” or minor weight-bearing areas of the patient and implanted into the prepared defect. Wagner used this technique for the first time in 1964. Hangody et al. and Bobic have been performing further developments since 1996 (Bobic, 1999;Hangody, et al., 1997;Hangody, et al., 1998). Hangody et al. implanted an average of 7 grafts per defect in 113 patients with mainly femoral defects from 1 to 8 cm2. After 3 years, the patients showed good joint functionality. Using the hospital for special surgery knee score system (HSSK) they determined scores from 91 (range 52 – 100), 91% of the patients showed pain relief and joint functionality after surgery. According to Hangody and co-authors in 10 years of clinical experience, this technique is ideally suited for osteochondral defects between 1 to 4 cm2 in patients less than 50 years old and without signs of osteoarthroses in another articulation (Hangody and Fules, 2003). The non-weight bearing patello-lateral periphery is usually selected as the donor site which later can be filled with fibrous cartilage. The treatment of large size defects with autografts is possible but implies the risk of donor-site morbidity. The surgical technique has recently been updated and is well documented (Hangody, et al., 2004). This technique, combined with osteotomy of the anterior cruciate ligament, showed good results (Ueblacker, et al., 2004). Lane and co-workers reported histological, chemical and biomechanical follow-up in a 6 months study of osteochondral defects filled with autologous plugs. The results showed bone restoration in all cases and a newly fibrous tissue filling the cartilage region. However, a gap at the interface between newly formed hyaline-like cartilage and host tissue was observed. The authors concluded that autogenous transplantation of osteochondral plugs is possible with integration of subchondral bone and preservation of chondral viability2 (85% at 6 months follow-up) although a cleft between the host cartilage and the differentiated cartilage was visible. An important observation in that study was that the stiffness of the graft from the donor site was not the same as that encountered in the recipient joint. They close the discussion by proposing the necessity to perform an experiment in which the stiffness of the donor site matches with the stiffness of the recipient site and to analyze its effect on osteochondral healing (Lane, et al., 2004).
The same surgical procedure as used in the autograft techniques can be employed to implant fresh or frozen donor osteochondral grafts. This technique is used for large, posttraumatic defects of the joints. The main advantages are that there is no limit for the number of plugs that can be used and that it can be prepared in any size. Therefore, this technique is used more frequently than the autologous one. The defects can heal to the host tissue and restore an articular surface (Buckwalter and Mankin, 1998). Fresh allografts and autografts allow a transplantation of viable chondrocytes. However, the usage of fresh allografts signifies not only logistic difficulties but also a risk of infection. The principal disadvantage of frozen allografts is the poor viability of the cartilage during long-term storage (20-30%). Some studies have been performed to increase this cell survival percentage by determining optimal banking conditions (temperature, cryopreservative substances and techniques, etc.) (Csönge, et al., 2002).
In a recent experimental study, the viability of press-fit preserved allografts to fresh autografts implanted into load-bearing and non load-bearing sites in mature sheep stifle joints was compared. The study showed in a 1-year follow-up that in all cases a better healing outcome was observed in the grafts implanted in load bearing regions. Histologically in all cases the line of demarcation between the graft and the host cartilage was indistinguishable. In a comparison of the viability (%viable/total cells) in the grafts implanted in the load bearing regions, the differences were not significant but they were slightly higher in the autografts compared with allografts (77.24% vs. 76.5%). In the non load-bearing regions 70% of viability was obtained for the autographs and 25% for the allografts (Gole, et al., 2004). After a comparison of all allografts versus all autografts implants, a viability of 55% vs. 77.5% was determined. By analyzing the results published by Gole et al. it can be concluded that the influence of the mechanical aspects on healing is higher as than the influence produced by biological conditions.
This technique refers to the usage of osteochondral grafts with tissues from donors of different specie than the receptor. In the majority of cases xenografts from pigs are cultivated in vitro and subsequently implanted in human patients. This technique has been developed as an alternative when there is a high risk for the patient (infection, morbidity of the donor site) to use autografts from non-load bearing regions of the patient or when a matching graft is not available due to the limited number of donors. Two possibilities have been investigated the usage of xeno-chondrocytes, which can be cultivated in vitro in an artificial matrix for the human cartilage or xeno-tissues, which are implanted into human defects. The usage of xeno-tissues has been studied experimentally using different animal models. The aim is commonly to reduce the inflammatory response and the delayed rejection of xeno-implants from donors to receptors. As part of the study of the xeno-implants, the process of xenogeneic graft rejection using pig chondrocytes implanted in mice was performed. The study identified a specific antigen and concluded that both native and in vitro transgenic cartilage from pigs could be used to treat cartilage defects in humans. The usage of xeno-transplantation and xeno-chondrocytes could be possible by controlling or eliminating such antigen (Costa, et al., 2003).
In a similar study, Stone and co-workers implanted pig cartilage in cynomolgus monkeys. They reduced the immune rejection of the xenografts by eliminating a specific pig antigen. Simultaneously they found that due to the elevated number of cells encountered in the synovial tissue, it showed a strong inflammatory response and concluded that the synovial tissue should be completely removed from cartilage xenografts (Stone, et al., 1998).
In an in vitro model, xeno-transplantation of pig chondrocytes to repair chondral defects in human cartilage was investigated. Here cultured cartilage implants were employed to treat defects of 2mm diameter and 2mm depth surgically created in femoral cartilage explants from donors. Histological analyses after 4, 8 and 12 weeks were performed. At the 12th week a hypercellular hyaline-like region rich in proteoglycans was observed, which showed very good bounding with the host cartilage. Formation of collagen type 1 and 2 in the newly differentiated tissue was detected. It was concluded that chondrocytes xenotransplantation could be used to repair defects in humans (Fuentes-Boquete, et al., 2004).
Some authors have reported that grafts improve their mechanical properties after a chemical process known as photooxidization. Additionally this process reduces the antigenetic response of the host tissues. The grafts are oxidized by immersion in an alkaline substance as methylene blue solution and are subsequently exposited to light with high frequency. In 1991 Nadler et al. presented results of a comparative in vivo study of osteochondral defects treated with different grafts in sheep. In their study a comparison between the healing outcome of autografts with and without photooxidation and xenografts after 12 and 18 months was performed. Histologically, fusion between host cartilage and the photooxidized grafts was found both times but it was not observed in non-photooxidized autografts and xenografts. Viability of the cartilage matrix was only detected in the oxidized grafts. In a similar study a comparison between the healing process between xenografts and autografts in sheep was performed and evaluated histologically after 6, 12 and 18 months. Parts of the xenografts were pre-treated with photo-oxidation. Although the xenografts were taken from bovine shoulder joint and transplanted into femoral joint defects no congruency3 problems were found. However, differences in the cartilage thickness were difficult to avoid. Such an experiment demonstrated fusion between photooxidized grafts and host cartilage. After 18 months, the hyaline cartilage structure was newly established and preserved at the interface photooxidized graft-host cartilage. Cartilage degradation and/or major gaps were observed in the control group as well as the autografts and xenografts without previous oxidization. Therefore the authors concluded that xenografts treated with photooxidization represent a good transplant possibility to treat osteochondral defects (Akens, et al., 2001).
In general, the usage of different resurfacing types of techniques such as autologous chondrocytes transplantation, in combination with treatment of axial malpositioning and ligament instabilities has showed to be an alternative for young patients with large defects (Steinwachs and Kreuz, 2003;von Rechenberg, et al., 2003).
The shape of the grafts appears to determine the success of healing. After long-term observation cylindrical grafts frequently show failures in the anchorage at the interface graft-host tissues. Therefore some investigators suggest that changes in the geometry of the graft could improve the healing outcome of osteochondral defects by altering mechanical conditions at the graft’s edges. In a comparative study, Rechenberg et al. analyzed the healing outcome of cylindrical and mushroom-structured osteochondral grafts. They found that the mushroom structured grafts differentiated into a less fibrous tissue, leading to an increased number of cells in the basal remodeling zones, a better restored subchondral bone and a minimal matrix degradation of the adjacent host cartilage after 6 months. The authors emphasize the importance of the subchondral bone for osteochondral bone survival and demonstrate the influence of the graft structure on the architecture of the subchondral bone and thereby the quality of the formed cartilage (von Rechenberg, et al., 2004).
In clinical practice combinations between techniques to stimulate repair and for transplantation are frequently found. One of the most frequently used is the implantation of cells in soft grafts and artificial matrix grafts. Autologous cultured chondrocytes and xeno-cultured chondrocytes are thereby employed. An alternative is the usage of differentiated chondrocytes from human stem cells, adult circulating blood cells, umbilical cord blood cells and more recently fetal circulating blood cells from rats (Naruse, et al., 2004). In vitro experiments have clearly established the necessary conditions to differentiate stem cells to chondrocytes (Barry, 2003;Barry and Murphy, 2004;Grande, et al., 2003;Hiraki, et al., 2001;Johnstone and Yoo, 1999;Luyten, 2004;Oreffo and Triffitt, 1999).
The development of techniques and products to treat articular defects as part of the topics considered by tissue engineering has obtained a remarkable importance during the last decades. However, a comparative study of diverse industrial tissue-engineered products in the world showed that after an initial boom economy declined drastically since 1995 (Lysaght and Hazlehurst, 2004). The study showed that this industrial sector could be on the borderline of a collapse. The authors compare data on all active firms in the field of tissue engineering in 1995, 1998, 2000 and 2002. The analyzed data include number of full time employees (FTEs), investment, number of firms created per year, annual spending and capital and the number of approved tissue-engineered (TE) products by the American Food and Drug Administration (FDA). The results show for example that the capital valuation of publicly traded firms was reducing from $ 1.9 billion in 1998 to $0.3 billion in 2002. At the end of the same year, twenty products had entered the FDA clinical trial. Four were approved but none of these are yet commercially successful.
An important conclusion from the literature, especially from the studies considering animal models, is that mechanical conditions appear to have a stronger influence on healing than biological or genetic ones. Therefore, although in this work spontaneous healing and the usage of defect fillings was evaluated without consideration of biological aspects, the predicted healing outcome after evaluation of the mechanical conditions on the repair process of osteochondral defects should be sufficient to evaluate a better treatment for a specific, individual situation.
Each technique has strengths and weaknesses. Independently of the technique used, the newly formed tissue usually consists of a fibrous or hyaline-like structure because although the new tissue shows morphological and biochemical similarities to the original hyaline cartilage, it does not possess a structure comparable with the host cartilage tissue. Hence, every surgical technique should be selected in accordance with specific conditions such as age and bone quality of the patient, and/or size of the defect to improve the mechanical properties of the tissues to be repaired. Perhaps the usage of a simulation tool, which is easy to handle and to implement, could help to determine more precisely how each of the involved parameters affects the final tissue quality. In this form, the determined mechanical conditions for an individual situation should be reproducible in vivo in order to obtain tissues capable to resist bearing loads without damage. This work mainly describes how such a simulation tool has been developed, evaluated and how it was applied to investigate the influence of mechanical conditions on healing.
Usually in vitro mechanical tests are used to determine mechanical properties of the cartilage. Due to the consideration of a biphasic behavior of the cartilage, it is necessary to determine the mechanical properties related with its fluid (permeability, aggregate modulus, etc.) as well as its solid phase (elastic modulus of Young’s, Poisson’s ratio, shear modulus, etc).
Before describing the more frequent tests used to determine mechanical material properties and the devices employed in these measurements, some basic engineering concepts shall be given as background information.
Uniaxial tensile or compressive loads are employed to calculate mechanical material properties. Uniaxial implies that stresses and strains are calculated in only one direction. After application of a tensile load (P) acting in an area (A) a deformation (ε) is expected. The deformation (ε) is calculated by relating the original length of the test specimen with the final deformed length. The tensile modulus of Young (E) is a measurement of the mechanical behavior of the specimen after load application.
Tensile Young’s module
Tensile loads produce not only an elongation in the axial direction (ε) but also a lateral contraction (εd) by forming a neck around of its neutral axis. Compressive loads would cause a shortening in the axial direction and a lateral expansion. Poisson’s ratio is a physical measurement that correlates the lateral strain (εd) with the axial strain (ε) by:
Where d0 is the original lateral dimension and d the deformed lateral dimension. Similarly L0 and L are the original and the deformed longitudinal dimension of the specimen, respectively. Poisson’s ratio magnitudes are usually tabulated for different materials in technical reports of public access. The Poisson’s ratio for isotropic materials varies between 0 and 0.5. A Poisson’s ratio of 0 corresponds to a material that is maximally compressible, and a Poisson’s ratio of 0.5 corresponds to an absolutely incompressible material. However, the determination of the Poisson’s ratio for biological tissues is still in permanent study. Cartilage, for example, has a Poisson’s ratio of 0.167 and bone has a Poisson’s ratio of 0.3.
Cartilage shall be considered as linear biphasic. The elastic compressive modulus (E), the Poisson’s ratio (ν), and the permeability (k) are the parameters that describe the mechanical behavior of the solid phase (extracellular matrix) in relation to the fluid phase (water). Permeability can be determined experimentally using Darcy’s law. Darcy’s experiment consisted of a fluid (water) going through a mechanism that allowed measuring the quantity of water absorbed by the solid content (sand). He established a relation between the volume of water, the size of the soil and the velocity by means of which the water flowed through the soil. The permeability (or perviousness) of rock (initially defined only for soils) is its capacity for transmitting a fluid. Degree of permeability depends on size and shape of the pores, size and shape of their interconnections, and the extent of the latter. It is measured by the rate at which a fluid of standard viscosity can move a given distance at a given interval of time. The unit of permeability is the Darcy4 (L2). The determination of the permeability in tissues such as cartilage is recent. A tissue with a high permeability has a low percentage of solids and vice versa. Cartilage e.g. has a higher permeability than the cancellous bone. Permeability is therefore normally closely related to the above mentioned material properties.
Materials are normally anisotropic. Yet in simplified representations a material can be modeled as orthotropic, transversally orthotropic or isotropic. An anisotropic material is characterised by possesses different material properties in all directions (En). An orthotropic material has different material properties in the principal three-dimensional directions E1, E2, and E3. A transversally orthotropic material has different material properties in a principal direction (E1) and in a transversal plane perpendicular to the direction of the load application (E2 = E3). An isotropic material has the same material properties in all directions (E1 = E2 = E3). In dependence of the material orientation assumed to represent a model with its material properties, several equations correlate mechanical parameters with each other.
In order to calculate these parameters (elastic deformation, elastic modulus of Young, etc.) cartilage samples were tested under compressive loads (Appleyard, et al., 2003;Boschetti, et al., 2004;Franz, et al., 2001;Wei and Messner, 1998). With the use of mechanical tests, it is intended to reproduce the mechanical boundary conditions activated during joint movements (axial, bending and torsional loads). Three different methods are frequently used to estimate the cartilage material properties: confined compression, unconfined compression and indentation, whose differences are based on the boundary conditions assumed during testing (Bae, et al., 2004;Goldsmith, et al., 1996;Korhonen, et al., 2002;Korhonen, et al., 2002;Lyyra, et al., 1999;Ming, et al., 1997). In correspondence to these boundary constraints, changes in the fluid direction and thereby in the material properties are expected. When a joint is in contact with any another joint, e.g. at the knee joint, and the fluids do not have perpendicular restrictions relative to the direction of the compressive active load, these boundary conditions are referred to as confined compression. In contrast, when the fluids inside the cartilage do not have movement restrictions in the radial direction and the fluids could move only in a perpendicular direction of the acting load, an unconfined compression constraint is defined. An indentation test measures the deformation at the cartilage surface after compressive load application using an appropriated device (indenter; see below).
Indentation is a mechanical test in which a device (indenter) applies a compressive load to the articular surface to be characterized. The indenter has strain gauges coupled to the indenter surface, which is able to measure deformations of the cartilage surface. The force (P), applied with the indenter and the contact area indenter-cartilage (A) are known parameters. The compressive stress during the test is calculated as the linear relation between these two parameters (P/A). Compressive stress versus deformation during the test can be plotted. The slope relating these two parameters represents the compressive elastic modulus of Young (E). The following equation correlates the different mechanical parameters involved during an indentation test through (k).
Using this equation and assuming different values for (ν) the parameter (k) can be calculated. The principal advantages of this mechanical test are that it does not require special specimen preparation, it is made in situ - thereby the boundary conditions are near to the ones encountered in the physiological situation - and it does not cause specimen damage. Stress controlled indentation testing of simple chondrocytes has been performed using a novel creep cytoindentation apparatus (CCA) (Koay, et al., 2003).
This mechanical test consists of the application of a compressive load on the specimen, in this case cartilage, which is fixed between two plates (Fig. 1.4a). The cartilage is then deformed in a direction parallel to the applied force (lateral). During load application the displacement of the deformed surface is then measured. The elastic modulus of Young (E) can be determined as the slope of the linear region in a plot of the history of load application versus the corresponding history of displacements. Using optical techniques it is possible to measure the lateral cartilage deformation. Knowing the lateral and the longitudinal deformation, the Poisson’s ratio can be determined. An unconfined compression test has been performed directly in chondrocyte cells using a novel system (a combination between a cell cultured nuclear microscopy and a laser indenter) (Leipzig and Athanasiou, 2005). An elastic modulus of 2.55 ± 0.85 KPa was measured using a linear elastic model. Similar values were found using a viscoelastic (2.47 ± 0.85 KPa) and a biphasic model (2.58 ± 0.57 KPa) to characterize the cartilage’s mechanical behavior.
This mechanical test is similar to the unconfined compression (Fig. 1.4b). But in this case the specimen is additionally constrained in the radial direction to the applied load. The specimen is placed in a chamber and a constant compressive load is applied to it. The lateral constrain avoids free lateral cartilage deformation developing a lateral pressure. The fluids inside the cartilage are then pressurized.
|Fig. 1.4: Schematic illustration of unconfined (a) and confined (b) compression test for cartilage samples (Boschetti, et al., 2004).|
Creep deformation occurs as the fluid is forced to flow from the tissue. Plotting a graph of the compressive stress versus the corresponding measured displacements, the slope shows the effect of the fluid pressurization, this parameter is known as aggregate modulus (HA). The aggregate modulus is calculated at creep equilibrium. The following equations correlate the aggregate modulus (HA) with the compressive modulus of Young (E) and shear modulus (μ) for an isotropic material.
Taking into account the pressurization of the fluid phase of the cartilage implies the consideration of another mechanical parameter of fluids: the permeability. Athanasiou and co-workers reported values of the aggregate modulus (HA) of 0.47 ± 0.15 MPa and k 1.4 ± 0.60 e-15 m4N-1S-1. The permeability is a parameter that relates the fluid velocity to the capacity of a material to absorb water.
Porosity is a mechanical parameter that relates the percentage of the solid phase and the fluid phase present in a biphasic material. This parameter can be evaluated from the water content, determined by the difference between the weights of hydrated and dried samples. Cartilage for example has a porosity of 0.8. That means that cartilage consist of 80% water and 20% solids (aggrecans and extracellular matrix).
Generally ultrasonography appears to be a technique frequently used over the last years to estimate joint degeneration or to measure mechanical properties, and it has been employed to determine normal physiological values of diverse musculoskeletal regions. By determining physiological values in healthy tissues, the attempt was made to standardize musculoskeletal ultrasound measurements and to avoid erroneous interpretation of normal echoes as pathological values. Ultrasonography values from muscles, cartilage and tendons were measured and averaged in healthy individuals and presented by Schmidt (Schmidt, et al., 2004). Damage of articular surfaces can be determined with ultrasonography as well (Nieminen, et al., 2004).
Low-intensity ultrasound has been employed to stimulate differentiation of bone tissues (Korstjens, et al., 2004;Sakurakichi, et al., 2004;Tis, et al., 2002), of human mesenchymal stem cell into chondrocytes (Ebisawa, et al., 2004) or to analyze its viability proliferation and matrix production (Zhang, et al., 2003). Other authors have demonstrated that low-intensity ultrasound can even influence cartilage repair (Cook, et al., 2001;Zhang, et al., 2002) and cultured chondrocytes in 3D arrays (Nishikori, et al., 2002). Ultrasound has been employed to analyze the state of repair after graft transplantation (Hjertquist and Lemperg, 1971) or to predict histological findings in regenerated cartilage (Hattori, et al., 2004).
In order to attenuate and to distribute uniformly the acting external loads, during joint functionality the fluids are moving physiologically with a slow and continuous velocity through the cartilage (Buckwalter and Mankin, 1998;Buckwalter and Mankin, 1998). Fluid velocity is directly related with the mechanical material properties of the cartilage by the magnitude of its permeability. Permeability was initially calculated experimentally for soils using the Darcy’s law (Equation 1.1, Equation 1.2). With the Darcy’s experiment the hydraulic conductivity (K) of a material can be calculated. Permeability is a parameter that depends on hydraulic conductivity (Equation 1.2). Since the Darcy experiment implies some technical complications to measure this cartilage parameter, it is frequently calculated in an indirect way using the finite element method. As input data the elastic modulus of Young, Poisson’s ratio, compressive strains, and/or aggregate modulus of the cartilage are required.
ΔV0/Δt flow rate (Rate of volume V in a time t)
Coefficient of permeability or hydraulic conductivity
gross cross sectional area of flow
total head (L)
length of flow path
With the equation 1.1 (Darcy’s hydraulic conductivity) can be calculated.
Knowing the permeability can be determined by equation 1.2
Where n =
g = gravitational acceleration
According to Forchheimer’s law, high flow velocities could reduce the effective permeability and consequently “chocking” pore fluid flow. As the fluid flow reduces, Forchheimer’s law approximates Darcy’s law (= 0).
Forchheimer’s law (permeability)
Permeability depending of porosity (n)
Void ratio (e) dependent on the porosity (n)
Porosity (n) depends on the strain (ε). Where n0 is the initial porosity (Equation 1.6)
is the volumetric flow rate of wetting liquid per unit area of the porous medium (the effective velocity of the wetting liquid)
is the fluid velocity
is the void ratio
is the wetting fluid volume in the medium
is the void volume in the medium
is the volume of grains of solid material in the medium
is the volume of trapped wetting liquid in the medium
is a “velocity coefficient”, which generally depends on the void ratio of the material
is the density of the fluid
is the specific weight of the wetting liquid
is the permeability of the fully saturated medium, which can dependent on the void ratio (e) and/or temperature θ
is the wetting liquid pore pressure
is the position
In Abaqus, a finite element solver, the permeability is defined as:
is the fully saturated permeability,
so that the Forchheimer’s law can be written as:
Remark: then have units of LT -1 . However, some authors use the hydraulic conductivity, which has units of L 2 (or Darcy). To use the saturated permeability in Abaqus, it needs to be multiplied by / . Where is the kinematic viscosity of the wetting liquid (the ratio of the liquid’s viscosity to its mass density).
Permeability (k) can be determined by numerical approximations using the finite element method. With the equations 1.4, 1.5 and 1.6 the problem is completely defined. Experimentally the elastic modulus of Young (E), the Poisson’s ratio ( ) and the compressive strains ( ) are determined from an unconfined test. Assuming porosity, void ratio is calculated from the equation 1.5 (Abaqus uses void ratio as input data). By inserting n from the equation 1.6 in the equation 1.4 an expression of permeability depending on strains can be determined.
The permeability is selected such that the strains measured in the unconfined test experiment are reproduced. When the equation is satisfied, the permeability of the tissue has been found. Once k is known, with the values obtained for fluid velocity ( ) and pore pressure ( ), the fluid mechanical parameters of the tissue are fully determined (Equations 1.7 and 1.8).
A theory for osteochondral repair and its computer simulation should take two concepts into account: bone remodeling and tissue differentiation. These principles and concepts describe the mechanical and biological conditions, which are active in an osteochondral repair process (Bagge, 2000;Scully, et al., 2001;Wang and Dumas, 2002).
The main aim of this project was the development and implementation of a tissue differentiation model, consistent with the histological findings, to evaluate the influence of mechanical conditions on osteochondral healing. Using linear monophasic material properties, the effect of the defect size, and cartilage thickness on healing was analyzed (Duda, et al., 2005). Subsequently, poroelastic non-linear biphasic materials were used to analyze the effect of joint curvature and the use of different defect filling stiffnesses in osteochondral repair.
The predictive tissue differentiation model was developed based on a theory for differentiation, which includes the bone remodeling concept and the acquired knowledge of tissue differentiation taken from osteochondral healing of animal experimentation. Prior to the development of a theory supporting the tissue differentiation model (see discussion), current models to simulate bone remodeling and tissue differentiation were investigated.
The bone-remodeling concept dates back to 1638 when Galileo Galilei established a relation between the bone geometry and the acting external forces. He understood that without an intelligent internal structure bones should possess an extraordinary size in order to transmit external forces in an optimal way (Nigg and Herzog, 1994). Subsequently, other authors who studied this internal structure currently named cancellous bone and its characteristics; mechanical behavior and remodeling capacity were explained.
The bone remodeling concept clarifies how cells in a tissue (bone or cartilage), in dependence of acting external loads and boundary conditions active during this process, are able to absorb or to produce another new tissue (Aro and Chao, 1993;Bagge, 2000;Kerner, et al., 1999;Tsili, 2000;Wang and Dumas, 2002). Some research groups have proposed tissue differentiation models with material properties (e.g. stiffness, permeability), taken from literature. These models have principally been used to study bone remodeling (Kerner, et al., 1999;Mikic and Carter, 1995) and fracture repair (Ament and Hofer, 2000;Claes and Heigele, 1999;Lacroix and Prendergast, 2002). Before this project, no differentiation model existed to simulate or to study osteochondral repair.
In 1884 Wolff found a similarity between the cancellous bone structure and orthogonal curves known from analytical representations of strains and stresses in a loaded beam (Fung, 1993). Bones were then represented by an internal configuration whose trajectories and form evidenced the existence of an implicit dependence of the applied external loads. In 1895 Roux discovered that such trajectories were the consequence of a dynamic functional adaptation of the internal bone structure reacting on externally applied loads. Bone tissue was able to change and to adapt its mechanical properties in dependence of current load situations (Fung, 1993;Huiskes, 2000;Huiskes, et al., 2000;Nigg and Herzog, 1994;van der Meulen and Huiskes, 2002). Thus, bone degrades its density distribution when external loads are reduced (zero gravity or immobility) or as consequence of a disease which destroys its internal original structure (e.g. osteoporosis) (Aloia, et al., 1987;Buckwalter and Mankin, 1998;Buckwalter, et al., 1994;Buckwalter and Brown, 2004). In this last case, a total destruction of cartilage has to be expected.
The determination of the mechanical cartilage properties is another topic of actual research. These are directly related to their remodeling capacity in the sense that the elastic modulus of Young is commonly selected to quantify the grade of remodeling or degradation of a tissue. To know the mechanical properties of the tissues is therefore fundamental.
In the following models this fundamental knowledge is used to explain how cells respond to changes in their mechanical environment stimulating biological processes such as growth or resorption. Histological analysis, in vivo and in vitro experimentation were and still are necessary to illustrate them. These models shall be introduced briefly.
Bones are able to adapt their internal structure in presence of mechanical stimuli generated after load application (Roux’s law). Different pathways for cell mechanotransduction have been identified to explain how this process could take place. The interpretation of these signals allows the development of tissue differentiation models, in which the mechanical cell environment is simulated and analyzed. For interpretation of this remodeling process Huiskes used the existence of an equilibrium state, originally proposed by Frost, in which cellular activity does not occur. An intact situation stays in the equilibrium state. Hence, equilibrium could be defined as a state in which the groups of cells responsible for growth or resorption are inactive and the balance between the acting external loads and the generated mechanical response causes the equilibrium. Remodeling is a complete cycle of cell activity in which after perturbations (prostheses, defects at bone, new boundary conditions) the equilibrium is newly reestablished. In this process different types of cells (osteoclasts and osteoblasts) are involved. This group of cells is called a basic multicellular unit (BMU). A BMU consists of about 10 osteoclasts and some hundreds osteoblasts. These groups of cells can be considered to be a mechanism of bone repair. Some authors believe that biochemical reactions are involved in the activation of these cells. Huiskes and his group propose that the BMU have some mechanosensors, which respond in the presence of external loads. Hence, deformations (a direct measurable consequence of the applied load) could be used as the mechanical stimulus to initiate a remodeling process. In other words, when an elastic, axial compressive load is applied to a tissue (for example bone), the derived compressive deformations act as mechanical stimuli to initiate remodeling.
Huiskes used this idea to develop a numerical model in which the knowledge of the equilibrium state of each bone region calculated in an intact bone situation was necessary. A stimulus for differentiation is activated as a reaction to the equilibrium alteration. The remodeling process stops when a tissue with the same mechanical properties as defined for an intact situation is completely reestablished in each affected region. The mandatory model of an intact bone must be calculated in order to know the strain magnitudes of the equilibrium state. After comparison of the straining of an affected and an intact situation, the necessary stimuli are calculated as a difference of these values. Huiskes schematized this process as a trilinear curve where maximal and minimal values of strains were used to define the range of change for each bone region. In this curve three points are defined to delimit the zones for growth, resorption, or equilibrium (dead zone) of each tissue as a function of a related stimulus (Huiskes, 1997;Huiskes, et al., 2000;Huiskes, et al., 1987) (Fig. 1.5). When the current strain falls below the value limiting the equilibrium zone, resorption sets in. In contrast, when the magnitude of the current strain is higher than the limits defined in the equilibrium zone, growth will occur. Following this schema, the elastic modulus of Young is continuously changed in an iterative process. With this model Huiskes calculated changes in the bone density (directly related to the elastic modulus of Young’s) that occur after implantation of a hip prostheses (Huiskes, 1990;Huiskes, 1997;Huiskes, et al., 1987;Kerner, et al., 1999;Mullender and Huiskes, 1995;Prendergast and Huiskes, 1995;Prendergast, et al., 1997;Vena, et al., 2000). This model was also applied to analyze the effect of stress shielding and how bone remodeling occurs after total hip replacement. In that study a special emphasis of the interface bone-implant was made.
|Fig. 1.5: Huiskes’s model. Cell activity is represented by a tri-linear curve. High values of strain energy density conduce to growth activity, low values of strain energy density conduce to resorption activity. The dead zone represents cellular inactivity (Huiskes, et al., 1987).|
Although Huiskes’s model showed a good correlation to the biological processes observed during and after a total hip implantation, this formulation was inappropriate to simulate fracture repair. Claes focused his work on the description of the mechanobiological process involved during fracture healing. After animal experimentation the total process from callus formation to cortical regeneration, the last state in a fracture repair, was documented and reconstructed through histological analysis at fixed points in time. In the Claes’s theory cells response to local strains and stresses, whose magnitude, position, and type could be determined. Claes modeled the geometry of the histological sections at the previously defined fixed points in time from his animal experimentation. Viscoelastic material properties taken from literature were used in these models. The finite element method was employed to determine mechanical strains and hydrostatic pressures for each tissue type (Fig. 1.6a).
|Fig. 1.6a:Claes’s model. Characteristic limit values of strain and stresses are defined for each tissue type to simulate diaphyseal fracture healing (cortical remodeling through enchondral ossification) (Claes and Heigele, 1999).|
The histological section showed how bone tissue was newly formed from an existent bone or cartilage region in an intramembranous ossification process. Cell differentiation for each tissue occurred in dependence of the acting strains or stresses whose magnitudes were previously determined with the computer model. Claes found that connective or fibrous tissue could be formed when hydrostatic pressures are higher than 0.15MPa and when the compressive strains are higher than 5% (Claes, et al., 1998;Claes, et al., 2002;Claes and Heigele, 1999).
Recently Claes’s group extended their model incorporating fuzzy logic5 (Shefelbine, et al., ). This method uses a combination of mechanical stimuli to simulate differentiation with series of fuzzy sets. Some rules (for unclear boundary conditions) need to be consecutively evaluated to describe a specific tissue type during healing (Fig. 1.6b). The rules are based on in vivo observations of the fracture healing process.
|Fig. 1.6b: Fuzzy logic in combination with mechanical stimulus applied to simulate healing of trabecular fractures (intramembranous ossification). Fixed rules including biological factors regulate building or resorption of different tissue types (Shefelbine, et al., ).|
Claes’s group demonstrated the possibility of using a general tissue differentiation model for fracture healing. In this form, the fuzzy logic model, which was initially employed to simulate diaphyseal fracture healing (Ament and Hofer, 2000), was further developed and applied to simulate metaphyseal fracture healing and bone remodeling. The first one refers to healing when the fracture occurs in the cortical bone. Cortical bone fractures heal by enchondral ossification: after a fracture, a callus replaces the formed haematoma. The callus is differentiated to fibrous and later to hyaline cartilage before it is replaced by cortical bone. The second one, metaphyseal fractures, compromises the trabecular bone. Trabecular bone fractures heal through intramembranous ossification: The bone is remodeled by activation of BMUs (basic multicellular units); osteoclasts and osteoblasts absorb and build new bone directly in dependence of biological and mechanical boundary conditions without cartilage formation. As a mechanical stimulus for differentiation they used octahedral strains in combination with hydrostatic strains.
Fuzzy logic is however a controversially discussed method because it anticipates the response of the healing process (or the system to be evaluated ) showing logically a high dependency on the fixed rules. However, the major contribution of Claes’s proposal is its intent to unify the fracture healing process for cortical and trabecular bone, allowing the analysis of fracture healing for different types of bones and different fracture localizations. Perhaps even more important is the point that his model allows the incorporation of biological factors observed during in vivo fracture healing such as different rates of vascularization and concentration of growth factors for the early and the late stages of healing.
Prendergast analyzed fracture repair under another perspective. Since following haematoma formation in the first stages of fracture repair only connective tissue is recognizable, which has a higher content of water, Prendergast proposed that cells react principally under the effect of this fluid phase. Fluid velocity and strains as a result of external loads should play an important role in the development of appropriate mechanical boundary conditions to initiate the repair process. Animal experimentation was used to define strain limits to start and to maintain cellular differentiation. During healing it was observed that micromovements at the bone surface stimulate bone repair. Using a biphasic, pore elastic model the mechanical shear strains and fluid velocities for each differentiated tissue during the repair process can be calculated. A mechanoregulation schema that combines these two mechanical parameters was then proposed to describe fracture healing (Fig. 1.7).
|Fig. 1.7:Prendergast’s model. Prendergast developed a mechanoregulation schema regulated by shear strains and fluid velocities defined for each tissue type (Lacroix and Prendergast, 2002).|
It was determined for example that a combination of high values of shear strains and fluid velocities cause distortion in the mesenchymal stem cells acting as stimulus to differentiate these cells to connective tissue. In a similar way cartilage is formed after a combination of low values of shear strains and slow fluid velocities (Lacroix and Prendergast, 2002;Prendergast, 1997;Prendergast, et al., 1997;Prendergast, et al., 1996).
Bailon-Plaza developed a predictive fracture healing hyperelastic model, in which growth factors were taken into account as an important parameter during repair. Plaza and co-workers proposed that growth factors are responsible for initiation and regulation of biological process of fracture healing. From in vivo studies the chondrogenic and osteogenic effect of two different growth factors were determined. The ability of these growth factors to regulate differentiation appears to be strongly related with the number of cells. A numerical implementation of this theory was used to simulate fracture repair. The model was able to predict and to demonstrate that the formation of osteogenic growth factors from osteoblastic cells and the duration of their production are necessary to obtain a full bone restoration after trauma. In her computer model dilatational and deviatory strains were used as stimuli to regulate the production of growth factors present principally in the first stages of fracture repair. The material properties used in her model were taken from literature (Bailon-Plaza, et al., 1999;Bailon-Plaza and van der Meulen, 2001).
Since the study of bone-joint mechanics and osteochondral healing implies a very complex biological process, the knowledge of biological aspects related to bone healing, studied by Lill and his group (Hepp, et al., 2003;Lill, et al., 2002;Lill, et al., 2001;Lill, et al., 2003), and osteochondral healing, studied by Bail and co-workers (Bail, et al., 2003), were used.
The works conducted by Lill’s group were basically employed to validate the model created in this project to analyze bone - joint mechanics (humerus project) and to perform a more realistic simulation of the bone specimens. In this project validation was made by comparing the measured stiffness in the mechanical test (compression and torsion) realized in human specimens to the mechanical stiffness calculated after simulation of these same tests in the same simulated bones. Following model validation, the influence of mechanical conditions on bone healing was evaluated through the analysis of straining of intact and fractured proximal humeri under physiological loads (in the rest of this project the corresponding sections are entitled bone-joint mechanics). The study in the proximal humerus realized by Lill’s group was selected because it presents well documented data (e.g. bone stiffness, implant stiffness), compares mechanical parameters of standard with recently developed implants to stabilize proximal humerus fractures and their precise measurements of the bone density distribution allows the development of a more realistic bone model.
The work realized by Bail was employed to validate the tissue differentiation model created to analyze the influence of mechanical conditions on osteochondral healing. Geometry, boundary conditions, histological and histomorphometrical analysis from the animal experimentation were used. As explained in the material and methods section, this validation was performed by comparing the histological and histomorphometric analysis to the simulated healing qualitative and quantitatively.
The work of Lill and co-workers is briefly introduced. In the biomechanics laboratory of the Musculoskeletal Research Center in Berlin, a previous study considering different bone qualities was realized (Lill, et al., 2002). Some parameters of humeral human bones were determined: stiffness, density distribution, trabecular bone orientation and volume bone fraction. Both intact and fractured human bones, stabilized with different medical devices, were mechanically tested and analyzed. The goal of Lill and his group was to determine favorable implant stiffness for fracture healing, taking into account the effect of the bone density distribution. Important implications for the treatment of fractures in osteoporotic patients were established.
Histomorphometrical, CT and MRI analysis were performed on 24 freshly harvested human cadaveric humeri. Histomorphometric analysis evaluated structural parameters (tissue volume to bone volume ratio, trabecular thickness), connectivity (number of nodes6, node to node length), and trabecular orientation (mean bone length). Median ages of 46 (34 – 46 years) and 69 (46 – 90 years) were registered for the male and the female group independently. In four horizontal levels sliced from the humeral head, five regions of interest were defined in each cutting plane: anterior, posterior, lateral, medial and central (in accordance with the standard coordinate system used for human bodies). With the information obtained (structural bone parameters, connectivity and trabecular orientation), a complete description of the bone quality of each specimen was determined. Different grades of bone strength were identified in the analyzed specimens: peak values of histomorphometric parameters at the cranial section decreasing caudally were found. This information was translated to the simulated specimens in the present project: localized differences of the bone quality could then be modeled.
The center of the trabecular structure connects the center of the gleaned cavity. A correlation between the structural parameters of the trabecular network, its localization and the estimated bone quality was determined: higher bone stiffness was found in regions of high trabecular density and elevated number of nodes compared with regions of low trabecular density. The project of Lill and his group aimed to obtain knowledge of distribution, microstructure, and quality of bone in the humeral head, which allows the remaining bone stock to be used effectively, even in elderly patients, with a minimally invasive approach and maximum mechanical stability (Hepp, et al., 2003;Lill, et al., 2002;Lill, et al., 2001;Lill, et al., 2003). According to Lill’s findings, bone quality appeared to be an important parameter to be evaluated. Therefore, the effect of the bone quality and its relation to the mechanical conditions (physiological loads) in intact and fractured bones during bone healing was then evaluated in this project.
In a subsequent study, 35 fresh human humeri were used to perform mechanical testing in order to determine in vitro characteristics such as stiffness of different conventional and new devices used to stabilize proximal humeri fractures. The implants tested under static and cycling loading included humerus T-Plate (HTP), the cross-screw osteosynthesis (CSO), the unreamed proximal humerus nail with spiral blade (UHN), the Synclaw Proximal Humerus Nail (Synclaw PHN) and the angle-stable Locking Compression Plate Proximal Humerus (LCP-PH). Three clinical load cases were evaluated: axial compression, torsion and varus bending. The results showed that the evaluated conventional devices (HPT, CSO, UHN, Synclaw PHN) presents a higher stiffness under static load than that encountered in the LCP-PH. Additionally, the torsional stiffness was strongly reduced with exception of the LCP-PH.
Lill and co-workers concluded that implants with low stiffness and corresponding elastic properties could minimize peak stresses, which could be related to the early loosening and failure of the interface implant-bone. Therefore low stiffness implants seem particularly suitable for fracture fixation in osteoporotic bones (Lill, et al., 2003).
The animal experiment realized by Bail and co-workers is briefly summarized (Bail, et al., 2003). Osteochondral defects of 6 mm diameter and 1.5 mm depth from the subchondral bone plate were created at the left femoral condyle in 18 Yucatan minipigs.
Under general anesthesia an osteochondral defect was created at the lateral surface of the trochelar groove of the left hind limb. The osteochondral defect was 6 mm in diameter and 1.5 mm in depth from the osteochondral junction (Fig. 2.2). To minimize soft tissue damage a sharp tube was pressed into the joint cartilage to the osteochondral border for guiding a 6 mm drill. A sleeve on the drill ensured the depth of the osteochondral defect. All wounds were sutured and covered with spray bandage. Animals received flunixin (Finadyne, Essex, Great Britain) as an analgesic for the first 7 days. Six animals were sacrificed after 4 weeks, nine after 6 weeks and the remainder after 12 weeks (Bail, et al., 2003).
In order to achieve a full load condition at the affected joint, an osteotomy at the contra lateral hind leg stabilized with a 8 to 10 hole DPC was created after surgery. To obtain a continuous load condition the animals were allowed to move freely after defect creation. The animals were on average 17.1 months old (10.5 to 30 months).
The following boundary conditions were reproduced in the model definition to perform osteochondral healing simulation: joint geometry and the defect size were modeled to resemble the mini-pig’s knee joint, and the total load was continuously applied from the initial defect situation for the rest of the healing process miming the in vivo load situation.
The animals walked on a compressive sensible platform. The measurements were made one day before surgery, three days after surgery and each week until the sacrifice of all animals. A complete gait cycle was defined as the enfolded step realized for a hind leg simultaneously with a foreleg. For each member 7 imprints were recorded (to minimize errors in the measurements). The weight of each animal was controlled during the realization of each gait analysis. After gait analyses an equivalent compressive load of 1.35 MPa was calculated from reaction forces of 1 body weight (BW) acting on an average area of 150 mm2 at the femoral condyle.
The histological sections were stained with Safranin-Orange van Kossa and Safranin-Light Green. This staining allows a clear identification of the formed tissue types. Safranin-Orange van Kossa stains the calcified tissues (e.g. subchondral bone) black and the non-calcified tissues (hyaline cartilage, fibrous cartilage and connective tissue) orange. Safranin-Light Green stains connective tissue and bone green, cartilage and cell nucleus are colored red. The mechanical tissue quality of the total group was slower than the corresponding of an intact situation.
The new formed tissues (hyaline and fibrous cartilage), the remodeled bone, and the remaining defect were quantified during healing by a histomorphometrical analysis (KS400 image analysis system, Zeiss, Germany). In three regions of interest (ROI), localized within the subchondral bone at the walls and at the basis of the defect, the structural orientation and the trabecular bone fraction were determined. For control purposes all histological and histomorphometrical analysis were made at the same regions of the contra lateral uninjured femur condyle. All statistical analyses (p<0.05) were carried out using a commercially available package (Statistical Package for Social Sciences, SPSS Inc., Chicago, USA) (Bail, et al., 2003). The trabecular bone fraction was increased indicating active remodeling regions.
Fraction of cancellous bone increased from 25% of the total area to 29.1 ± 10.6% at 4 weeks, subsequently to 33.0 ± 13.9 % at 6 weeks and to 33.2 ± 8.7 % at 12 weeks. Hyaline cartilage was increased from initially 0% to 8.1 ± 6.7 % at 4 weeks, 17.0 ± 12.5 % at 6 weeks, and 33.1 ± 25.4 % at 12 weeks. The fibrous tissue filling the defect decreased from 47.7 ± 14.2 % at 4 weeks to 35.2 ± 13.2 % at 6 weeks, and to 24.2 ± 18.7 % at 12 weeks. The defect region was reduced from initially 75% of the total area to 15.2 ± 7.71 % at 4 weeks, to 14.8 ± 9.0 % at 6 weeks, and to 7.1 ± 2.9 % at 12 weeks. During healing it was found that the trabecular bone fraction increases indicating active remodeling regions (Bail, et al., 2003).
Additionally, Bail et al. reported a complete macroscopic and microscopic description of the repair process. New techniques such as a reproducible immunohistochemical color protocol for cartilage and bone repair analyses were developed.
From Bail’s animal experiment the following data was used:
The study of spontaneous repair of osteochondral defects using animals for experimentation were realized in line with a doctoral thesis in veterinary medicine entitled “Histologische, immunhistologische und histomorphometrische Untersuchungen der Wirkung von systemisch applizierten Wachtumshormon auf einen osteochondralen Knorpeldefekt am Yucatan-Minischwein” (histological, immunohistological and histomorphometrical analysis of the effect of systematical applied growth factors on an osteochondral defect in Yucatan minipigs) (Klein, 2001).
In this thesis all sections related to osteochondral defect healing have the title “Galileo project”. Historically, Galileo Galilei is considered to be a pioneer in the field of biomechanics. He not only used the term “mechanics” for the first time to relate acting forces, systems and its response, but also observed and applied such concepts to biological organisms. Galileo supposed for the first time a relation between bone and its mechanical boundary conditions. His contribution to biomechanics was the basis for further development of bone remodeling and tissue differentiation (a more general formulation). Such concepts have been refined, further developed and are recently called mechanobiology (van der Meulen and Huiskes, 2002). Therefore, the name “Galileo” was used to identify the study of the mechanical aspects of the biological process involved in osteochondral healing.
Before studying osteochondral healing the mechanical behavior of the bone-joint region was studied in a large model. The straining of the intact and fractured proximal humeri under physiological-like loading conditions was determined. The importance of tissue quality, for the straining of bone in intact and osteosynthetic stabilized proximal humeri, was demonstrated as a result. This project aimed to determine the influence of the mechanical conditions on osteochondral healing. To achieve this goal the development of a tissue differentiation model, which is able to predict healing with appropriate tissue quantification, was necessary. After validation by a comparison with the outcome from animal experimentation this model was used to answer some of the most frequently reported clinical questions, such as which is the maximum defect size that still produces tissues with a good mechanical quality, why osteochondral defects occur predominately on convex joint surfaces and, in the case of defect fillings, how stiff should a biomaterial be in order to guarantee a long-term joint functionality. The tissue differentiation model developed in this poroject demanded the proposal of a theory supporting its usage and implementation. Additionally, this project explored for the first time the usage of in vivo data to determine mechanical parameters (minimum principal strains, fluid velocity, pore pressure etc.) of the different tissue types during healing, some of which (e.g. fluid related material properties) are difficult to determine without the use of numerical tools.
1 Burden of disease: refers to the combination of the incidence/prevalence, impact (in terms of quality of life and disability) and cost of musculoskeletal conditions. Defined by the Bone and Joint Decade 2000-2010 Council.
2 Viability: Refers to the capacity of cells/tissues to live, to develop or to germinate under favorable conditions. Capable of living outside of the uterus. Used for a fetus of a newborn: quality or state of being viable. That is, the ability to live, grow and develop.
3 Congruency problems refers to negative biological process (necrosis of the tissues in contact) derived from inefficient assembly, after geometrical differences, between the donor and the receptor site of the graft.
4 Named after H. Darcy (1803-1858), one Darcy corresponds to the volume of a liquid with a dynamic viscosity of 1 centipoise flowing through a porous material with a cross section of 1 cm2 in 1 second when undergoing a pressure drop of 1 atm/cm. This unit was used in hydrology and civil engineering to measure the permeability of porous materials.
5 Fuzzy logic: Is a superset of boolean logic dealing with the concept of partial truth. Whereas classical logic holds that everything can be expressed in binary terms (1:yes; 0:no), fuzzy logic replace these simple boolean truth values with degrees of truth. This allows for values between 0 and 1, the concept of “maybe”. It allows partial membership in a set (i.e. elements can belong to different sets simultaneously). The method can use logical connectors to describe unclear events or events that cannot be represented with unique true values. A membership function is normally required to adjudicate the grade of certainty of the variable to be evaluated. In this form unclear events can be treated with IF/THEN rules for final association of a deterministic success.
6 Nodes: trabecular intersections at the cancellous bone.
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