Non-Metallic Biomaterials for Bone Substitutes and Resorbable Biomaterials in Orthopaedics

Transcription

1 University of Patras Faculty of Medicine Department of Medical Physics National Technical University of Athens Department of Mechanical Engineering Interdepartmental Program of Postgraduate Studies in Biomedical Engineering National Technical University of Athens Department of Electrical Engineering and Computer Science Non-Metallic Biomaterials for Bone Substitutes and Resorbable Biomaterials in Orthopaedics Master of Science Thesis Baciu Diana Elena National Technical University of Athens Department of Electrical Engineering and Computer Science Supervisor: Prof. D.KOUTSOURIS Athens, 2004

2 Supervisor Professor: D. Koutsouris, Professor National Technical University of Athens Department of Electrical Engineering and Computer Science Examination Board: D. Koutsouris, Professor National Technical University of Athens Department of Electrical Engineering and Computer Science K. Nikita, Associate Professor National Technical University of Athens Department of Electrical Engineering and Computer Science P.Tsanakas, Professor National Technical University of Athens Department of Electrical Engineering and Computer Science

3 Acknowledgments Firstly, I wish to express my deepest gratitude to my supervisor, Professor Dimitris Koutsouris for his continuous availability, his valuable guidance and advices in completing this work and also for his trust on me during this study. I feel also indebted to Professor K. Nikita for her cooperation, guidance in the realization of this work, for her help, patience and great attitude. I would like to thank also the Prof. P.Tsanakas for his valuable contribution in the realization of this work, for his patience and well coming attitude. It was a pleasure for me collaborating with Dr. D.Scarpalezos, to whom I would like to express my warmest thanks for his significant contribution, for the knowledge he offered me concerning orthopedic biomaterials and the meaning conversations. I would like to thank the firms Biomet and Vamvas for all the useful stuff and for their advices. I would like to thank the Professors of the Department of Electrical Engineering and Computer Science in the National Technical University of Athens for the knowledge offered, their help and support during this study. I feel the need to express my sincere gratitude and special thanks to Professor Nicolas Pallikarakis for his moral support and especially for making my study in Greece possible. Finally I wish to express the most loving thanks to my parents and my husband, who were beside me all along, encouraging and understanding me.

4 Table of Contents Abstract Chapter I. Introduction in biomaterials... 1 I. 1. Biomaterials-general description...2 I. 2. Orthopedic biomaterials. General...3 Chapter II. Bone and non-metallic biomaterials...12 II. 1. Bone-anatomy and mechanics...13 Compact bone...14 II.1.1. Structure...14 II.1.2. Elastic behaviour...15 II Anisotropy...16 II Mineralisation...16 II.1.3. Non elastic behaviour...17 II Viscoelastic behaviour...17 II Plasticity...17 II Fracture...17 II.1.4. The affect of age...19 Cancellous bone...19 II.1.1. Structure...19 II.1.2. Mechanical properties...20 II.1.5. Adaptive behaviour II The sensor...22 II The controlling variable...23 II Modelling the process...23 II.1.6. Non mechanical variables...24 II. 2. Non-metallic biomaterials for bone substitutes...24 II. 3. Modern non-metallic biomaterials for bone substitutes - manufacture, properties, clinical applications and results...30 II.3.1. Calcibon-synthetic bone substitute...31

5 II.3.2. Collapat II-Bone substitute biomaterial based on collagen and hydroxyapatite...38 II.3.3. Endobon...40 II.3.4. Palamed G-the new generation of bone cements...44 II. 4. Bone cement and modern cementing technique...47 Chapter III. Bioresorbable biomaterials in orthopaedics. General III. 1. Modern bioresorbable biomaterials in orthopaedics -manufacture, properties, clinical applications and results...69 III.1.1. Topkin-degradable wound cover...69 III.1.2. Inion HEXALON-the first coloured biodegradable ACL/PCL Screw...73 III.1.3. ReUnite Resorbable orthopedic plating system-made exclusively from LactoSorb (resorbable copolymer)...77 III.1.4. Gentle Threads-resorbable orthopedic fixation system-made exclusively from LactoSorb...78 III.1.5. ReUnite Resorbable orthopedic fixation system made exclusively from LactoSorb...79 III.1.6. Inion Trinion-meniscus screw with controlled insertion...83 III.1.7. Inion OTPS-biodegradable pins...85 III.1.8. Inion OTPS-the first complete biodegradable fixation system for the ankle...86 Chapter IV. Conclusion and discussions...89 Bibliography

6 Abstract Thanks to recent advances in science and engineering, the field of biomaterials stands poised to increase the effectiveness and longevity of established devices as well as to provide new options to biomedical engineers who work at designing future products. From its beginning, the field of bioengineering has focused on providing the best artificial devices - hearing aids, artificial limbs and other prostheses - to replace body parts that are missing, broken, or dysfunctional. Regeneration of body parts requires a biomaterial with a structure, components and chemical signals that allow the body s tissue cells to recognize, respond to, and remodel the material without rejecting it as foreign. Bone, cartilage and the major load bearing joints of the body all suffer degenerative changes with age and trauma. This area of research focus will seek solutions to the problems of osteoporosis, the fixation of implants in bone and the replacement of damaged bone and cartilage. This will be achieved through the development of non-metallic biomaterials and resorbable biomaterials that provide appropriate load bearing characteristics and the potential to interact suitably with the biology. The non-metallic materials for bone substitutes serve as scaffolds and may have modified surfaces to encourage natural tissue growth or the ability to be seeded with the hosts own cells before implantation. This will have applications for both bone and cartilage substitute materials. These bone substitutes biomaterials are the second-most implanted of all materials. Resorbable biomaterials, on the other hand, gradually disappear from the body as a result of hydrolysis, because are made from molecules similar to those in the human body, which resorb while the tissue is healing. This eliminates the need for a second surgery. The goal of this thesis is to explain the usefulness of these biomaterials in medical applications and especially in orthopaedics, focusing on the latest acquisitions. The first chapter makes an introduction in biomaterials, with emphasis on orthopaedic biomaterials. The second chapter contains information about: (1) the bone characteristics (anatomy and mechanics), in order to understand the basis for tissue engineered therapies and how damaged bones heal, (2) the non-metallic biomaterials (polymers, biodegradable polymers, ceramics and composites) for bone substitutes, giving examples of modern biomaterials used today and (3) the principles involved in the Modern Cementing Technique. The third chapter is a review of the chemistry of the polymers used in bioresorbable biomaterials, including synthesis and degradation, describe how properties can be controlled by proper synthetic controls such as copolymer composition, highlight special requirements for processing and handling, and presents in detail some of the commercial resorbable biomaterials.

7 Chapter I. Introduction in Biomaterials Chapter I Introduction in Biomaterials Motivated by a need for custom-made materials for specific medical applications, material scientists, chemist, chemical engineers, and researchers in other disciplines have turned their attention to creating high performance biomaterials. Biomaterials science examines the mechanical, physical and chemical properties of materials as well as the complex host responses to introduced bulk material, material surface and biomaterial applications. Biomaterials science has been officially defined as the study and knowledge of the interactions between living and non-living materials, and biomaterial as a material intended to interface with biological systems to evaluate, treat, augment or replace any tissue, organ or function of the body. The development of modern biomaterials is related to the development of modern medicine and new materials. Stainless steel and cobalt chromium alloys were the first materials successfully used inside the body for fracture fixation. In the early 1960s, Sir John Charnley made the first attempt to link together a stainless steel hip prosthesis and high-density polyethylene with metachrylate bone cement. This can be considered the beginning of modern orthopedics, in which the development of better materials plays a central role. In the late 1960s, the excellence of titanium was discovered in medicine. At that time, some materials began to be classified as biomaterials. Various materials (polymers, ceramics, composites and metals with improved properties) and applications (orthopedics, vascular and heart surgery, etc.) have been developed since then. Today, there are a great number of different professions dealing with the problems associated with biomaterials, and the cross-scientific approach is essential. Tissue engineering is an emerging field that aims to regenerate natural tissues and create new tissues using biological cells, biomaterials, biotechnology, and clinical medicine. There are international organizations, which give recommendations and standards for the manufacturing and testing of biomaterials (ISO= International Standards Organization, ASTM= American Society for Testing and Materials). There are also national organizations that supervise biomaterial applications in human use. One of the best-known and most demanding control organizations is FDA (Food and Drug Administration of USA). It must be pointed that FDA does not regulate the materials used in medical devices, but rather the devices themselves.[1] Lawsuits against medical device manufacturers, restructuring of FDA approval procedures, patient expectations, and the health care reform movement are changing the future of the medical device community and shaping the direction of biomaterials research. For example, lawsuits have prompted long-term material suppliers and device manufacturers to refuse the use of their products in medical applications. As a result, new materials and suppliers will be required to meet FDA standards. 1

8 Chapter I. Introduction in Biomaterials I. 1. Biomaterials-general description Biomaterials are defined as materials that are used in clinical medicine in direct contact with human cells and tissues such as blood, cornea, bone, and subcutaneous tissue. Are natural or man-made (polymers, ceramics, composites and metals with improved properties), that comprises whole or part of a living structure or biomedical device which performs, augments, or replaces a natural function. [2], [3] Biocompatibility, which is a central factor, has been officially defined as the ability of a material to perform with an appropriate host response in a specific application. The biocompatibility of a medical implant will be influenced by a number of factors, including the toxicity of the materials employed, the form and design of the implant, the skill of the surgeon inserting the device, the dynamics or movement of the device in situ, the resistance of the device to chemical or structural degradation (biostability), and the nature of the reactions that occur at the biological interface. These factors vary significantly depending on whether the implant is deployed, for example, in soft tissue, hard tissue, or the cardiovascular system to the extent that "biocompatibility may have to be uniquely defined for each application." Biomaterial-tissue interactions are also described by the term biocompatibility, which is the ability of a material to elicit an appropriate biological response in a given biological application. This implies that any material placed into the body will not be inert and will interact with tissues in a dynamic way, altering both the material and the tissues around it. The biological response to a material is critically dependent on three factors: (1) the composition-structure of the material, (2) characteristics of the host site, and (3) the functional demands on the material. Thus, a material cannot be characterized as biocompatible or not biocompatible until its host site and function are defined. Biomaterial-tissue interactions are relevant to a wide variety of disease states and treatments. Traditionally, this area has focused on materials, which can support or replace lost human tissue, for example in dentistry (filling materials/implants/material-tooth bonding), orthopedics (femoral implants), and vascular biology (arterial stents). Biomaterials have also been used in more esthetic applications such as maxillofacial prostheses to restore facial contours after surgical treatment for cancer, or in dentistry to restore tooth esthetics after trauma or dental disease. More recently, biomaterials-tissue interactions have been involved in an area called tissue engineering or bioengineering. Roughly defined, tissue engineering is the manipulation of developmental or wound-healing processes to repair or replace lost tissue. In this area, biomaterials have been used as permanent or resorbable scaffolds (for cells or other critical molecules), drug-delivery devices, and barrier materials. For example, cartilage cells may be seeded onto a resorbable scaffold, implanted, and allowed to support cartilage development. Or pancreatic islet cells may be encapsulated in a material, which allows transport of only small molecules and nutrients, but not immune cells. In dentistry, collagenous or polytetrafluoroethylene membranes have been used as barrier to prevent growth of epithelial elements and allow regeneration of bone around teeth or implants. In any case, biocompatibility issues are critical to the success of these treatment strategies. [4] 2

9 Chapter I. Introduction in Biomaterials Biomaterials are: Orthopaedics joint replacements (hip, knee), bone cements, bone defect fillers, fracture fixation plates, and artificial tendons and ligaments. Cardiovascular vascular grafts, heart valves, pacemakers, artificial heart and ventricular assist device components, stents, balloons, and blood substitutes. Ophthalmics contact lenses, corneal implants and artificial corneas, and intraocular lenses. Other applications dental implants, cochlear implants, tissue screws and tacks, burn and wound dressings and artificial skin, tissue adhesives and sealants, drug-delivery systems, matrices for cell encapsulation and tissue engineering, and sutures. Some Commonly Used Biomaterials: Material Silicone rubber Dacron Cellulose Poly (methyl methacrylate) Polyurethanes Hydrogels Stainless steel Titanium Alumina Hydroxyapatite Collagen (reprocessed) Applications Catheters, tubing Vascular grafts Dialysis membranes Intraocular lenses, bone cement Catheters, pacemaker leads Ophthalmologic devices, drug delivery Orthopaedic devices, stents Orthopaedic and dental devices Orthopaedic and dental devices Orthopaedic and dental devices Ophthalmologic applications, wound dressings The above biomaterials are in widespread clinical use. [5] I. 2. Orthopaedic biomaterials. General. The arena of orthopaedic medicine is widely researched and has led to improved surgical procedures in areas such as bone grafting, joint replacement, and limb supports. Finding suitable biomaterials for a variety of applications is a key aspect of orthopedic research. Orthopaedics is the branch of surgery, which is concerned with the preservation, and restoration of function of the skeleton, its articulations and associated structures. The name orthopaedics comes from the first applications of the modern field in the past few hundred years: treating musculo-skeletal deformities in children. The field of orthopaedics dates back to ancient times and has an extremely illustrious history. 3

10 Chapter I. Introduction in Biomaterials Orthopaedics is currently one of the largest medical industries and deals with deformities, injuries, and pathology of the bones, joints, ligaments, tendons, and muscles. In addition, the field is as technologically advanced as any other in medicine. Orthopaedic subspecialties (material applications): 1. Adult reconstruction (joint replacements) 2. Orthopedic oncology (limb salvage materials) 3. Spine (fusion, fracture fixation, curve correction systems) 4. Pediatrics (fracture fixation) 5. Hand (fracture fixation) 6. Foot (fracture fixation) 7. Sports - Upper extremities (shoulder, elbow reconstruction) - Lower extremities (knee, ankle reconstruction) 8. Trauma (fracture fixation) Naturally occurring orthopedic biomaterials (tissues): 1. Bone - Cancellous - Cortical 2. Cartilage - Articular 3. Ligament 4. Tendon 5. Meniscus 6. Intervertebral disc - Nucleus pulpousus - Annulus fibrousus Fact 1: Bone is only tissue that heals without scar formation Fact 2: Some orthopedic tissues have little or no healing potential making tissue replacement necessary 1. Cartilage (artificial joints) 2. Some parts of the meniscus (removal) 3. Intercapsular ligaments (ACL, PCL) (ligament grafts) Fact 3: All orthopedic tissues have pathological conditions where they won t heal Fact 4: Most tissues will regenerate if damage is in very young (making tissue engineering an important area of research) Graft materials (allograft and autograft): 1. Bone - Cancellous - Cortical 2. Cartilage - Articular 3. Ligament 4. Tendon 5. Meniscus 4

11 Chapter I. Introduction in Biomaterials Allograft issues: 1. Sterilization - Diseases (HIV, Hepatitis B and C) - Mechanical properties (properties degrade with sterilization) 2. Immunogenic response Types of traditional artificial orthopedic biomaterials: 1. Plaster of Paris or fiberglass casts 2. Metals-characterized by metallic bonds: - 316L stainless steel - Co-Cr-Mo alloy (cast) - Co-Cr-W-Ni alloy (wrought) - Titanium - Ti-6Al-4V Alloy Issues: - Note, Ni and Cr are toxic ions in first 3 metals - The more biocompatible alloys are more expensive - Most compatible is not the hardest 3. Ceramics-inorganic compounds with ionic and covalent bonding: - Aluminum oxide Issues: -Hard and biocompatible but brittle -Osteophilic 4. Polymers-long chain molecules of repeating units: - Bone cement (PMMA) Issues: - Exothermic - Toxic effects of monomer - Shrinkage with polymerization - Grout, not adhesive - UHMWPE (polyethylene) Issues: - Lipid absorption - Wear particles - Low coefficient of friction and creep resistant 5. Resorbables (make hardware removal unnecessary) - PLA or PGA or Co-polymers Issues: - Adequate strength - Strength over time -Degradation without adverse biologic reaction. 6. Others- glass (non-crystalline solids) and carbon. Basic Atomic Interactions: Bonds between atoms Ionic - electrostatic bonding Covalent-electron sharing Metallic-electron fluid or gas Hydrogen-ionic interactions between covalently bonded atoms 5

12 Chapter I. Introduction in Biomaterials Van Der Waals-shifting interactions between atoms Crystals Organized, repeating 3-D pattern of molecules or atoms - Closely packed structure Stress and Strain: Stresses are forces applied over areas Strain is a dimensional change due to an applied stress Axial stress and strain Tension +, compression Hooke s law- stress and strain proportional Young s modulus Shear stress and strain Poisson s ratio- lateral distension for axial load Material Properties: Material failure Yield stress Ductile and brittle failure Plastic deformation Ultimate stress and strength Some fail in shear, compression and tension Fatigue- failure under cyclic conditions though well below yield stress Creep- time dependent extension Stress relaxation Toughness Energy absorption to failure Material properties Consistent numbers not always available Variation in runs, machines, locations Structures generally a laminate composite Properties may be function of fabrication process or post-processing Measurement of key properties such as stress, Young s modulus, strength, and Poisson s ratio Surface Properties: In General Surfaces are uniquely reactive Surfaces are different from the bulk Surfaces are readily contaminated Surface material/ structure is mobile Can change depending on environment Surface structures or properties Roughness Chemistry or molecules Inhomogenous surfaces Crystalline or disordered Hydrophobicity (wettability) Contact angle 6

13 Chapter I. Introduction in Biomaterials Surface Measurements: Contact angle ESCA -Electron Spectroscopy for Chemical Analysis (XPS) Element identification and bonding state Auger Electron Spectroscopy SIMS (Secondary Ion Mass Spectrometry) Element ID, Low concentrations, Proteins FTIR- ATR (Fourier Transform Infra Red) Chemistry and Structure Orientation STM- Scanning Tunneling Microscopy SEM (Scanning Electron Microscopy) AFM (Atomic Force Microscopy)[6], [7], [8]. Component design and analysis: Advanced finite element analysis 3-D parametric computer-aided-design modeling Probabilistic mechanics and reability analysis Reability-based design optimization. [9] Fracture Fixation Systems are: nails plates screws external fixators Ilizarov systems external casting pins wires cables. [6] Examples of fracture fixation systems: Anchor System The Corifix Ligament Used for the replacement/repair of ACLs and PCLs using the hamstring tendons. Offers high pullout strength in both cyclic and static loading. Titanium construction for maximum MRI compatibility. 7

14 Chapter I. Introduction in Biomaterials GS Interference Screw Manufactured from titanium alloy for optimum strength and biocompatibility. Accept as standard 3.5mm hex driver for positive screw placement. Supplied sterile. 16 sizes. The Actor Ligament Graft reinforcement device suitable for hamstrings or bone patella tendon bone. Ideal for use when poor hamstring graft has been harvested. Fixed via Interference Screws. ACL or PCL applications Radial Head Implant Machined from Cobalt Chrome Dual Coated hydroxyapatite on porous metal Used when Radial Head has been fractured and is displaced or unstable Restores axial stiffness of forearm to normal 8

15 Chapter I. Introduction in Biomaterials The Halder Humeral Nail Simple insertion without exposure of the fracture site. Retrograde insertion no damage to Rotator cuff. 30% stronger proximal fixation Lower profile Distal Plate. Easier Extraction Method. The Rigidyne External Fixator The Rigidyne External Fixator combines dynamisation, rigidity and versatility. The unique telescoping motion combines dynamic axial loading with rigidity. Adaptors allow the Rigidyne to be used as a hybrid incorporating ring fixators or pelvic bars. Contains pins and accessories, which allow the body of the fixator to be converted to a brace, type drill. Corifix Trauma Systems (Osteosynthesis Products) A cost effective range of trauma products, which incorporate high standards of manufacture and materials to ensure excellent clinical results in the most demanding operative situations. A comprehensive range of stainless steel screws and plates, which are available either sterile or non-sterile. [10] 9

16 Chapter I. Introduction in Biomaterials Mechanical properties of orthopaedic biomaterials: Typical properties of materials associated with fracture (or impending fracture) fixation. Variations in failure stress and fatigue stress for metals arise from how the material was formed (i.e. cast, cold worked, host forged, etc.), the surface condition, surface treatment and other factors. [6] Material Modulus (GPa) Poisson s Ratio Failure stress (MPa) Fatigue Stress (MPa at 10 6 ) Stainless steel (AlSl316) Titanium alloy (Ti-6Al-4V) Cobalt-chrome alloy (Co-Cr-Mo) PMMA (tension) (compression) Cortical bone (tension) (compression) Cancellouse bone (compression) (porous structure) PMMA-bone interface 7-10(tension) (shear) PMMA-metal interface 5-10(tension) 5-8(shear) 2.0 UHMWPE 1 Fibrous tissue Requirements for replacement materials: 1. Must re-establish mechanical function with healing or replacement - Structural and mechanical criteria - Strength - Stiffness - Fatigue life - Wear - Wear debris 2. Must not adversely affect biological environment - Biocompatibility criteria - No inflammatory response - Inflammation causes lymphocyte invasion, vascular occlusion, tissue necrosis, fibrous tissue formation, prosthesis / implant loosening - Not carcinogenic - Not toxic - Not mutagenic - No immunogenic response - Must not overload or underload surrounding tissues - Overloaded tissues will die - Underloaded tissues will resorb or atrophy - Unloaded tissues will not heal as rapidly and may have compromised tissue organization and microstructure 10

17 Chapter I. Introduction in Biomaterials - Must be appropriate size and not disturb surrounding tissues and blood supply, particularly those necessary for tissue healing - Must not allow too much or too little micromotion (relative motion at bone-bone or bone-implant interface) 3. Must not be adversely affected by biological environment - Non-corrosive (metals) - Stress - Galvanic - Crevice - Lipid absorption (polymers) causing property degradation 4. Manufacturing and use - Ductility (plates must be contoured during use) - Machinability 5. Cost 6. Requisite size (determined by mechanical properties) versus available space 7. Sterilizable 8. Size of material (i.e. wear debris) biocompatibility doesn t fix things. [7] This chapter was a briefly description of the biomaterials with emphasis on orthopaedic biomaterials. 11

18 Chapter II. Bone and Non-Metallic Biomaterials Chapter II Bone and Non-Metallic Biomaterials Bones defects occur in a variety of clinical situations like cancers or other diseases, that can destroy them, and many people are born with missing or deformed bones. Their reconstruction to provide mechanical integrity to the skeleton is indispensable in the patient s rehabilitation. Bones usually can regenerate themselves to repair defects up to a certain size, but if the defect is too large and severe, the suitable biomaterials have to be applied to facilitate the bone repair. The study of bone structure and mechanics is the first step in orthopedic biomaterial science. Bone is a complex tissue consisting of sponge-like cancellous bone and solid compact bone in varying proportions, depending on its location in the skeleton. Is constituted of nanosize blade-like crystals of hydroxyapatite(ha) grown in intimate contact with collagen fibres. The several components of the bones are: minerals to give them hardness; proteins to give them strength; blood vessels to nourish them; and special cells that build and remodel them. These special cells are osteoblasts, osteoclasts and osteocytes. Osteoblasts are the cells presents in the bones, which actually build the extracellular matrix and regulate its mineralisation. Cells named as osteoclasts are able to resorb fully mineralised bone as they are equipped with a variety of enzymes, which lower the local ph to values between 3 and 4. Osteocytes are the principal (they account for about 90% of all cells in the adult skeleton) cells present in adult bones, and their special construction may actually orchestrate the spatial and temporal recruitment of the cells that form and resorb bone. Together, they form a team that grows and remodels bones throughout life, as you grow taller, stronger, heavier, and older. The development of non-metallic materials (polymers, biodegradable polymers, ceramics and composites) for bone substitutes called also scaffolds, is very useful in bone healing, because they may have modified surfaces to encourage natural tissue growth or the ability to be seeded with host own cells before implantation. Tissue engineers can use the osteoblasts to grow new bones. They place these bonegrower cells in biomaterial scaffolding with the mineral component of bone. The cells use this structure for support while they produce the proteins and minerals to grow new bones. Placing growth factors in key areas of the scaffold helps shape the bone growth. In some cases, the scaffold is placed right on the bone defect in the patient, and new bone tissue grows into the scaffolding. Tissue engineers hope to be able to design a bone to match the shape of an individual patient. Computers will help by layering cells and biomaterials in two dimensions at a time, building towards the complex three dimensional structure of a real bone. 12

19 Chapter II. Bone and Non-Metallic Biomaterials II. 1. Bone-anatomy and mechanics The bones of the skeleton play an important part in the support and movement of the body, and also provide mechanical protection for vital organs. These functions are possible because the bones are both stiff and strong, as is the bone, which is a major component of the structure of the bones. Bones and bone material are not homogeneous but have a clearly defined structure. In the typical long bone, such as the femur, the central portion of the shaft consists of a cylinder of compact cortical bone, but the ends are filled with a spongy open textured cancellous bone consisting of a three-dimensional lattice made of bony trabeculae (Figure 1). Articular cartilage Spongy bone Trabecula Compact bone Periosteum Nutrient artery Collagen fibres Intramedullary cavity Line of epiphyseal fusion Macro Concentric lamella (3-7μm) Micro to ultra Fig. 1 Hierarchical levels of structural organization in a human long bone (femur) Osteocyte Outer circumferential lamellae Neighbouring lamellae having differ fiber orientations Haversian canal Fig. 2 The structure of compact bone: each Haversian system may contain up to 30 lamellae, and have a diameter of up to 0.1 mm 13

20 Chapter II. Bone and Non-Metallic Biomaterials Fig. 3 Diagram of a small section of a long bone. M: the side next to the marrow cavity; P: the side next to the periosteum. Three differently oriented test specimens are shown, all loaded along their long axes. L: longitudinal; T-tangentially; R: radial. Compact bone II Structure. Fig.1 shows a diagrammatic representation of a typical long bone. In young humans and cattle, and some other mammals, the structure of the compact bone consists of concentric layers of bone material about 5μm thick. This is lamellar (sometimes called fibrolamellar or plexiform) bone. In humans this is replaced during growth, the process starting with the formation of cylindrical cavities around the blood vessels (Figure 2). Specialized cells called osteoclasts carry out the resorbtion. The result of resorbtion is to provide a well-defined cylindrical cavity aligned with the long axis of the shaft of the bone. The first stage of replacing this cylindrical cavity (filled with tissue fluid) by bone is the laying down of a demarcation cement line at the surface of the cavity. There follows a process whereby osteoblasts lay down concentric layers or lamellae of bony material 3-7 μm thick. Surrounding the central blood vessel, there is a cylindrical cavity, which may also contain such things as nerves and lymphatics. The resulting elongated cylindrical bony structure is called a Haversian System (in the UK) or a secondary osteon (US). Between the lamellae of bone there are cells called osteocytes in lacunae, which interconnect with each other and ultimately with the central hole via a system of canaliculi. The health of these cells is vital to the bone; osteocyte death has severe consequences and will lead to the resorbtion of the bone by osteoclasts. The Haversian System will consist of 20 or 30 layers, resulting in the structure 0.1mm to 0.2mm in diameter and considerably greater in length. The long axis is aligned with the long axis of the bone. The lamellae consist of layers of fibrous protein called collagen. The collagen is made up of small constituent molecules called tropocollagen, which 14

21 Chapter II. Bone and Non-Metallic Biomaterials contain a helix of three protein chains. These are secreted by cells and organized extracellularly by being joined end to end to form long thin fibrils, which are then cross-linked laterally. This will look rather like a wall with bricks and mortar. Within the "mortar" region of the resulting collagen fibril, minerals are laid down. These are predominantly a form of calcium phosphate called hydroxyapatite, whose unit cell contains Ca10(P04)6(0H)2. This mineral occurs as small crystals in the form of needles or plates, which interact with each other so that if all the organic collagen material is removed we have something, which still looks rather like the bone. The crystals themselves are stiff and brittle with poor impact resistance being typical ceramics. The collagen fibres are much more extensible, and the combination of collagen and apatite gives a composite with properties different from and better than those of the constituents. The overall composition of the bone is inorganic material (mainly apatite) about 50%; organic material (mainly collagen) about 40%; water about 10%. Within each lamella the collagen fibres run predominantly parallel both to each other and the cylindrical surfaces of the lamella, i.e. they are spirally wound. The angle or pitch is relatively constant for each lamella, but varies between neighboring lamella. The angle may be almost at right angles to the long axis or at greater pitch. A small number of fibres run between the lamellae of a Haversian System. The outer surface of the Haversian System is demarcated from the neighbour by the cement line which few fibres cross. The Haversian Systems are approximately cylindrical and it is impossible to pack cylinders together to fill all the available space. Hence in the mature human bone there will be packing between the Haversian Systems, which consists of the remnants of older resorbed systems, or the original fibrolamellar bone. Outside the packed Haversian System there will be a few continuous bony lamellae, which go all the way round the shaft of the femur. II Elastic behaviour There are two general methods for investigating the elastic properties of bone: (a) Ultrasonic velocity -the velocity of sound in a medium (v) is given by: V = E /ρ, where E is the Young s modulus and ρ is the density of the medium (typically Kg.m-³) Ultrasonic testing is rather difficult and there is considerable argument on whether the methods are valid even in principle. The major problem is the inhomogeneity of the bone, which also contains voids. It is possible in theory to use wet and, indeed living, specimens, but for reasons of simplicity tests have been carried out mainly on dry specimens. (b) Mechanical test can be made on excised and shaped specimen of bone. Constant strain rate tensile or compressive tests on these specimens to strains of 1% or less suggest the material behaves as linear elastic solid. The slope of the stress strain curve (the tensile or compressive Young's modulus) depends on the direction in which the specimen has been cut, the degree of mineralisation, and whether the bone is wet or dry. 15

22 Chapter II. Bone and Non-Metallic Biomaterials II Anisotropy Anisotropy is apparent in the structure of the bones and this is further reflected in the mechanical properties. The results are simplest to understand if the specimens are cut longitudinally, tangentially or radially (Fig.3). The longitudinal tensile modulus is typically around 20GPa. The other two moduli are smaller, the tangential modulus being around 15GPa and the radial modulus about 13GPa. The other moduli and Poisson' ratio would also expect to show directional difference, but there is insufficient reliable data to establish this definitively. The shear modulus of specimens cut longitudinally is of the order of 3GPa and Poisson's ratio appears to be around 0.3 to 0.4, but there is a very wide range of values. Table 1.Some Physiological Proprieties of Three bone tissues: Property Antler Femur Bulla Work of fracture/j m-² 6,190 1, Bending strength/mpa Young modulus/gpa 7,4 13,5 13,3 Mineral content/%ash Density/10³ Kg m-³ 1,86 2,06 2,47 Table 1 shows the work of fracture for antler, femur and bulla, and it is quite clear that the value of this variable is strongly influenced by the mineral content of bone. II Mineralisation Mineralisation can be characterized in several different ways: The most usual is to quote the density of the bone, because the apatite is more dense than collagen and highly mineralised bones will therefore have a greater density. Alternatively the ash content can be quoted; the bones are dried at high temperature to calcine the organic material and drive off the free water. The normal range of variation of mineralisation in bones is relatively small, but tests on a wide range of bones from the same species suggest that there is a linear relationship between the tensile modulus and strength and the ash content. Suspending bones in weak acids can vary the degree of mineralisation. A new equilibrium level of mineralisation is produced, and this is reduced as the ph falls. The results of such test confirm that the more highly mineralized the bone, the higher the modulus. It is not clear what changing the ph does to the collagen, which is a protein and unlike all proteins contains groups, which are acid or basic 16

23 Chapter II. Bone and Non-Metallic Biomaterials depending on the ambient ph. Care must, therefore, be exercised in interpreting the results of such tests. Bones of different species or different types can be used to explore the effect of mineralisation. Table 1 gives the results of three different types of bone. Antler is lightly mineralised and bulla of whale is highly mineralised. This is apparent in both the ash content and the density in Table 1. There is a parallel increase in the modulus of elasticity. II Non elastic behaviour II Viscoelastic behaviour Viscoelastic behaviour is apparent as stress relaxation and creep, although the magnitude of these two viscoelastic manifestations is generally rather small. Strain rate dependence is also slight unless a very wide range rates are used. McElhaney carried out tests on bone over a strain rate range of 10 3 to 1.5 x 10 ³ and found a threefold increase in stiffness. It is really only at very high impact rates that the strain rate dependence is of major importance. There is considerable discussion about the structural basis of viscoelastic behaviour. Collagen fibers themselves are viscoelastic and show marked stress relaxation, although the mineral crystals are entirely elastic. It is not clear how the collagen and apatite composite will behave and it has been suggested that viscoelastic behaviour is largely associated with the cement lines around the Haversian System; the cement is certainly less highly mineralised and more likely to undergo stress relaxation, but there is currently no consensus but lots of argument. II Plasticity Carrying out a tensile test on a bone, which has been kept moist during the preparation and testing and has edges, which have no serious cracks or scratches, shows that bone undergoes considerable plastic flow. A quite different appearance occurs if the specimen is dry where sudden catastrophic failure occurs before the plastic region is entered. Once again there is considerable discussion about the structural basis of plasticity but it does seem likely that displacement of neighboring Haversian Systems at the cement line plays a major role in plastic deformation. II Fracture A single increasing loading or deformation of the bone can produce failure. The results of such test show that the tensile strength, the yield strain and the strain of fracture are all anisotropic (Table 2). The compressive strength is generally greater than the tensile strength. 17

24 Chapter II. Bone and Non-Metallic Biomaterials Table 2. Failure properties of Compact bone. Species Man Cow Cow Histology Direction of loading relative to long axis HS HS FL Parallel Normal Parallel Normal Parallel Normal Tensile Strength Yield strain " Ultimate strain Compressive Strength Yield strain Ultimate strain Failure tests can also be carried under three or four point bending, and the results generally show greater strength than do tensile tests. This is probably because in bending the strains are not uniformly distributed across the bone section, and the most highly strained parts of the specimen undergo plastic deformation, thus redistributing the stresses and strains in the specimen. Controlled failure tests can also be carried out using wide specimens with edge notches or cuts. The geometry allows slow crack propagation across the specimen, and more information can be obtained about the mechanics of fracture than can be from simple tension tests. Work in this area is just beginning and interesting results are now starting to emerge. Bone tends to be rather stronger than one might expect. It probably relates to the composite structure of the material. In a propagating crack there is a high stress concentration at the tip of the crack, because of its sharpness. If a crack propagates through an Haversian System it will meet a softer and weaker cement line before the next Haversian System. The crack can propagate alon g the cement line more easily than it can propagate through the new Haversian System. As a consequence the tip of the crack will be blunted and the stress concentration much reduced. This will also occur within the Haversian System between the lamellae and of course within the lamellae where there is collagen and apatite. This method of crack blunting is widespread in biological materials and also occurs in man-made composites. Some evidence for the incremental nature of crack propagation can be obtained by listening to bone undergoing failure (acoustic emission). As the bone is increasingly loaded there is an increase in the number of these failure events, which can be heard. 18

25 Chapter II. Bone and Non-Metallic Biomaterials Before cracks can propagate they have to start and tests specimens have to be prepared very carefully so that there are no initial cracks or scratches. Bone is notch sensitive and the present of such imperfections of the surface will very much reduce the strength. Bone itself contains holes both as the Haversian Canal and as the canaliculae, which might be expected to acts as stress raisers. They do not appear to do so. Repeated loading to stress less than the tensile strength will result in failure after a large number of cycles - a process called fatigue failure. In general the smaller the magnitudes of the repetitive stress relative to the tensile strength, the greater the number of cycles required to produce failure. There is often a linear relationship between the logarithm, of the stress value and the number of cycles to produce failure, but there is extraordinary variability between the reports of different workers. It is somewhat difficult to relate fatigue failure in excised bone specimens to what happens in the body. There is a well-known fracture of the metatarsal head, which occurs after repetitive loading during marching. This "March" Fracture is a well-known clinical condition and appears to be the only substantiated example of fatigue fracture occurring in the body. The failure properties of bone can be described either in terms of the failure strength as above or in terms of the toughness of the material. The toughness relates to the amount of energy absorbed by the material before it fails. The mechanical work is stored in the material as strain energy. Dry bone is commonly brittle, but wet bone is often rather tough, and stores a considerable amount of energy. A stronger bone, which has a slightly greater failure stress, may well be very much tougher because of the added extension in plastic flow, which is possible. II The affect of age The age dependence of the tensile properties of human femoral cortical bonethe stiffness remains relatively constant, but there is a fall in strength with increasing age. This may have severe consequences, because the older people are more unstable and may have impaired neuromuscular systems. One consequence of this is the increased number of fractures in the elderly, particularly fractures of the neck of the femur. Cancellous bone Cancellous bone is dense, apparently solid, stiff and strong. It has a much lighter open structure and occurs as padding in the cavity within the compact bone. II Structure Cancellous bone has a rather complicated and variable structure. The simplest description would be that it consists of apparently randomly oriented cylinders of bone about 0.1mm in diameter or plates of about the same thickness. These cylinders 19

26 Chapter II. Bone and Non-Metallic Biomaterials or plates extend for about lmm before dividing or connecting with another. This results in a three dimensional meshwork of bone with space between the cylinders or plates of bone. The space is usually occupied by marrow, although in birds it may be air. On a larger scale the cancellous bone appear to be random and isotropic or more ordered and anisotropic. The isotropic random type typically appears inside bones well away from the loaded surface, acting as padding. The bone which has clear orientations, rather like trajectories or buttresses occurring on buildings, are generally found in areas of load such as beneath the articular surface or at such sites as the trochanter. It is clear from the description of the structure of cancellous bone that it must be very porous. The greatest value of porosity is not far short of 100%, when the marrow is interspersed by a few spars of bone. The lower limit is more difficult to set, but bones with a porosity of about 50% can be described either as dense cancellous bone or as compact bones with rather a lot of holes in it. The difference between cancellous and compact bone is also reflected in the apparent density. Compact bone has a density of typically 1800 to 2000kg.m-³; cancellous bone will have a density from about 200 to 1000 to 2000kg.m- 3. II Mechanical properties Almost all the mechanical tests reported on cancellous bone have been made in uniaxial compression. This is because it is extremely difficult to obtain the elongated specimen required for tensile tests, and there is considerable difficulty in holding the specimens if you do have them. Compression tests are very much easier, but they require considerable care. During most tests there is some interaction between the grips or loading platterns and the specimens. In compression tests this will take the form of friction between the loading plattern and the bone specimen, and unless the specimen is reasonably long compared to its transverse dimensions, this can seriously disturb the results. However, making specimens long and thin while overcoming this problem introduces another one. Such specimens tend to buckle when loaded rather than undergo the compression we need. During the initial stage of compression, the stress-strain relationships are linear, but yield and an extensive region of plastic behaviour where considerable strains occur with little change in stress follow this; this is followed by a phase when the bone specimen once again becomes stiffer. The behaviour can be closely related to the structural changes in the specimen during loading. The post-yield behaviour at constant stress is reduced by buckling and collapse of the constituent" spars and plates of the bone. When all these are collapsed the boney material is compacted and is no longer porous and this accounts for the stiffer region before total failure. There is a variety of data in the literature but that of Carter and Hayes is probably the most reliable. Their results shows that the log /log plots of both strength (S) and density (D) and compressive modulus (E) and density are both linear. 20

27 Chapter II. Bone and Non-Metallic Biomaterials E=kD³ S=k¹ D² Carter and Hayes data are unusual in the wide range of densities used and over a more restricted range the relationship may appear to be linear, and this is reported in several papers. More recent work and tensile strengths and modulus suggest that they obey the same parallel as compression and that the constants k and k 1 were nearly the same. It is quite difficult to understand this relationship. Considerable work has been done on the behaviour of open cell foams, which can be applied to bone. The results strongly suggest that both S and E should be proportional to D 2. It certainly matches the strength behaviour of bone, but the discrepancy in a modulus behaviour remains inexplicable. There is also information on the shear strength, which has been given as between 5 and 7 MPA. The value of the strength would be expected to depend on the porosity, which is generally not given. II Adaptive behaviour Bones give the impression of being inert, but this is far from being true. The material in the bone is continually exchanged with similar material in the blood. Calcium plays an important role in cellular activity, particularly of active cells of muscle and nerves. The maintenance of the necessary calcium level for this purpose will override the maintenance of calcium in the bone. When the diet is seriously deficient in calcium or there is an exceptional need, as in pregnancy, the bones may well lose their calcium content and undergo some demineralisation. Similarly there is a continual change and turnover of the constituents of the collagen, although this is very much slower. At a higher-level bone is continuously being remodelled, with osteoclasts removing bone and osteoblasts replacing it. Bone also responds to functional loads. Long disuse by being in bed or being subjected to weightlessness in spacecraft leads to demineralisation and disorganisation of the structure of bone. This can have important consequence for patients, who have undergone serious illness, e.g. have fractured their hips and have been maintained in bed without exercise for long periods. Similarly, increasing the functional load on a bone will lead to increased mineralisation and structural re-organisation. The effects of load have serious consequences when placing internal implants such as artificial hips. If the design is such that some areas of bone are not loaded - a phenomenon caused stress shielding - they will undergo resorption and may lead to the loosening of the implant. Bone responds differently to tensile and compressive loading, this is most clearly apparent in orthodontic tooth movement during which forces are applied to the teeth by wires or similar means. The tooth is held in the bony tooth socket by a periodontal membrane, which consists of collagen fibres which bridge, the gap between the tooth and the bone. When forces are applied to the crown of the tooth, some parts of the membrane and hence the bones are under tensile force, some parts under compressive. The bone, which is subject to 21

28 Chapter II. Bone and Non-Metallic Biomaterials tensile force, responds by laying down new bone, but the bone subjected to compressive force resorbs. In trying to explain how bone maintains its normal structure despite its being in an essential dynamic equilibrium, and how it responds to abnormal situations is necessary to explain how the mechanical state of the bone can be sensed. Also is recommended to try and identify some variable, which the bone uses as the optimum setting and it would be helpful to model the process in some mathematical way. II The sensor There is considerable discussion on the way bone senses the mechanical loads or deformations to which it is subject. It has been proposed that osteocytes are the primary sensor, and calculations, making assumptions about Poisson s ratio and about modulus of water, suggest that they could be subjected to transient high pressures when the bone is deformed. The major criticism of the theory is that bone remodelling is essentially a surface phenomenon and the osteocytes are locked deeply inside the Haversian Systems. The osteocytes interconnect with each other and presumably the surface via processes in the canalliculae. Recent work has also shown that when bone is subjected to repetitive loading, changes do occur in the osteocytes, e.g. an increased take up of tritiated uridine and incorporation into RNA. Bone, and many other collagenous tissues, has strange electrical properties such that when deformed or loaded they generate potential differences. There is considerable discussion on how these stress-generated potentials (SGP) are produced. Piezoelectric behaviour has been postulated. Many crystals, which do not have a centre of symmetry, will generate a potential when strained. This is not true of apatite, but it is true of collagen, and it is conceivable that the collagen/apatite composite will also show some piezoelectric effects. Streaming potentials have also been suggested as a cause of the SGP. Bone contains channels full of solutions of electrolytes. The surface of the bone, like most biological surfaces, is charged, and this charge will attract some of the ions in the electrolytes solution, and bind them. If the electrolyte solution is forced to flow, the bound ions do not take part in this flow, and the resulting flow of charges produces and electrical current and potential difference. Most of the work, which has been done on SGBs, has been done on moist bone, maintained at something like 95% relative humidity. There are results, which shows that if is bended a cantilever of bone, a potential difference develops across it with the side tension being positive in respect to the side in compression. The peak potential occurs during and immediately the deformation is complete and gradually falls to zero, although this rate of fall may be strongly influenced by the measuring technique. The potentials are quite large and have been reported of being of the order of 5-10 mv. The main problem in discussing SGBs is how to relate what happens in moist bone at 95% relative humidity to what happens in fully saturated bone permeated by an ionic solution in the body. No really satisfactory answer has yet 22

29 Chapter II. Bone and Non-Metallic Biomaterials been given to the problem. Nevertheless there is considerable evidence that electrical phenomena can alter remodelling and fracture healing. II The controlling variable There is no general agreement on what is the controlling variable. Strain has been suggested and this seems to be the best established. Studies of surface strains during normal function (obtained by sticking strain gauges to bones invivo) show that for many studies the normal peak physiological strains in adult bone whose main function was load bearing fall in the range % - this is impressive unanimity. Other candidates have been local stresses and the amount of energy stored per unit volume when the bone is strained, but these are rather less well substantiated. II Modelling the process Bone responds to an altered mechanical environment by one or two sorts of changes. The gross structure may change with additional bony trabecullae, which also may be reoriented. The material properties may also change producing alterations in the interrelated porosity, density and modulus. Modelling has largely concentrated on the latter group of variables. The simplest linear elastic adaptive theory was developed by Cowan who suggested that there was a physiological strain (e0) to which the remodelling aimed. He related the Young's modulus (E) the actual strain occurring (e) and this ideal strain by the simplest linear equation: (de/dt) = A(e-e0) t being time and A a constant. He similarly described the external surface remodelling of bone by: dx/dt = B(e-e0) where X was a characteristic surface coordinate perpendicular to the surface and B a constant. The equations will predict that where the strain difference is zero there is no remodelling, that if e is greater than e0 there will be an increase in stiffness, and if e is less than e0 there will be a reduction in stiffness. This is a very interesting theory. The As and Bs and the strains are all second order tensors and the theory is developed in a fairly general tensor form. The only other reasonably developed theory is that of Fyhrie and Carter who produced a linear model connecting the change in density of the bone material with the stored energy produced by loads or deformations. 23

30 Chapter II. Bone and Non-Metallic Biomaterials II Non mechanical variables When discussing the physical properties of bones and bone, it is important to remember that non-mechanical variables may also be very important. The thermal property is one example because orthopaedic implants are usually fixed into bones using some heat generating luting material, e.g. acrylic cement. If the surface of the adjacent bone heats up too much the osteocytes and other vital cells will be damaged as can the collagen of the bone material itself. There have been several models, mainly FEM models, which seek to investigate this heat, transfer process. [11] II. 2. Non-metallic biomaterials for bone substitutes (scaffolds) The importance of orthopaedic substitutes and healing therapies is underscored by the following news story: The World Health Authority has decreed that will be the Bone and Joint Decade, and this is now being supported by the United Nations. The rationale for this is that joint diseases account for half of all chronic conditions in people over 65; back pain is the second leading cause of sick leave; and osteoporotic fractures have doubled in the last decade, so that 40% of all women over 50 will eventually suffer from one. It is estimated that 25% of health expenditure in developing countries will be spent on traumarelated care by the end of the decade, and many children are deprived of normal development by crippling diseases and deformities." To address the need for non-metallic biomaterials for bone substitutes, current clinical therapies include: 1. Autografting involves harvesting a bone from one location in the patient s body and transplanting it into another part of the same patient. Autografting is considered the gold standard. An example of one the most commonly performed bone autografting procedures is for use in spinal fusion. In a fusion procedure, bone graft from the patient's hip is implanted in disc spaces between spinal vertebrae or along the back of the spine (Figure 1). The grafted bone fuses the vertebrae together over several months. The benefit from transplanting an autogenous tissue is obvious: immunogenicity is not an issue. Autografting, however, has several associated problems including the additional surgical costs for the harvesting procedure, and infection and pain at the harvesting site. For example, harvesting an ileac crest graft (i.e., the protruding bony section of the patient's hip) can cost between $1000 to $9,000/procedure for the harvesting operation and the additional hospital stay. The morbidity at the harvest site can be tremendous with problems such as pain, infection, and blood loss requiring blood transfusion adding the associated risks of transfusion reaction and blood borne infection. 24

31 Chapter II. Bone and Non-Metallic Biomaterials Fig. 1. Autografting procedure to repair collapsed disc. 2. Allografting involves harvesting and processing bone from a cadaver then transplanting it to the patient. Allogenic implants are acellular and are less successful than autografts for reasons attributed to immuogenicity and the absence of viable cells that become osteoblasts. Another disadvantage of allografting is concern with transmitted disease. 3. Man-made materials, including metals, plastics, and ceramics represent an important percent of bone substitutes. Cancellous bone graft has a high content of osteogenic stem cells but lacks stability and must be used with stable internal fixation. High morbidity of the donor site is yet another drawback of this procedure. Cortical bone graft has a low content of osteogenic stem cells, poor vitality and serves rather as a bridge-forming splint. Immune defence, potential for HIV infection and contamination are disadvantages of the use of allo- and heterografts. Bone defects are also treated using bone segment transport, vascularised bone graft, bone marrow aspirates and demineralised bone powder. Alternative treatment modalites involve growth factors, cytokines, gene therapy, and tissue engineering and bone graft substitutes. Bone has some regenerative capacity and therefore many tissue-engineering efforts focus on implantation of biomaterials and factors that will augment the natural repair process in large defects that would not normally heal. [12] An ideal bone substitute it should: be porous with interconnected pores of adequate size; allow for the ingrowth of capillaries, perivascular tissues and osteo-progenitor cells; attract mesenchymal stem cells from the surrounding area; promote their differentiation into osteoblasts; be bioresorbable-biodegradable, complex calcium ions from the body fluids, be radiolucent and mechanically strong. The latter feature is much dependent on the intended application of the substitute. The fact that substitutes expand the volume of autogenic graft, serve as scaffolds, enhance the stability of fixation and protect against the infiltration of soft tissue, which promotes graft consolidation is advantageous. [2] A scaffold provides cells with a three-dimensional matrix upon which to adhere, proliferate, and produce matrix and may be engineered to serve as a tool to create more complex tissue structures. Scaffolds can function as delivery vehicles for cells or they can perform a tissue inductive role such that, when implanted, cells from host tissue are encouraged to migrate into it and form functional tissue. The development of novel scaffolds for tissue-engineering applications has been a very active area of investigation. 25

32 Chapter II. Bone and Non-Metallic Biomaterials For a scaffold to support cell viability and tissue development, it must be porous to allow for effective transport of nutrients and waste products as well as being biocompatible with the host tissue. It should eventually degrade as tissue matrix is produced, by hydrolytic or enzymatic mechanisms, leaving nontoxic degradation products that can be eliminated from the body. In addition, monomers, excipients, cross-linking agents, solvents, and other residual chemicals that can potentially leach out of the scaffold postimplantation must exhibit minimal toxicity. Finally, the scaffold must have the correct physical and mechanical properties for the given application. The scaffold must be able to withstand physiologic loading until sufficient tissue regeneration occurs. [13] The following are functional and mechanical properties required for bone formation: strength formability toughness osteoinductivity controlled degradation inflammatory response Osteoinduction is a chemical process in which molecules within the bone scaffold change the patient's cells into bone forming cells. Osteoconduction is a physical effect by which the use of a scaffold generates the growth of matrix components from the host tissue. It is ideal for it is biocompatible and forms a favorable three-dimensional matrix for human osteoblast cells to adhere and spread, associating the advantage of collagen osteoinduction to the superior bioactivity and osteoconduction of HA. Non-metallic biomaterials for bone substitutes serves as scaffolds materials for making matrices for bone tissue engineering. They can be: -of inorganic origin; -of organic origin. The first include synthetic ceramics such as hydroxyapatite, tricalcium phosphate, carbonated hydroxyapatite, calcium sulphate, coral and ceramic cements. The second group comprises bioresorbable native and synthetic polymers. Composites based on polymers and ceramics provide another option. Synthetic Polymers, both organic and inorganic, are used in a wide variety of biomedical applications. The polymers can be biodegradable or nondegradable. Commonly used synthetic materials are poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and their copolymers, poly(lactic-co-glycolic acid) (PLGA). These biodegradable polyesters have a long history in medical applications and are amenable to many fabrication and processing techniques. A biodegradable scaffold is used to provide support for bone-producing cells during the healing process. Once the cells have replaced the missing bone, the scaffold dissolves away, and only the new bone is left behind. These have demonstrated the ability to support cell attachment, proliferation, and matrix production for a variety of cell types relevant for musculoskeletal tissue engineering, including chondrocytes, osteoblasts, and mesenchymal stem cells. Synthetic polymers enable precise control over the physiochemical properties such as degradation rate, porosity, microstructure, and mechanical properties. 26

33 Chapter II. Bone and Non-Metallic Biomaterials These FDA-approved polymers are currently used as suture materials, but are also being examined for uses such as bone, skin and liver substitutes. These polymers are broken down in the body hydrolytically to produce lactic acid and glycolic acid, respectively. Fig.1. A cross-section of a PTFE Other biodegradable polymers currently being studied for tissue engineering applications include polycaprolactone, polyanhydrides, and polyphosphazenes. Polymethylmethacrylate (PMMA), polytetrafluoroethylene (PTFE), and PMMA/polyhydroxyethylmethacrylate (PHEMA) may be described as alloplastic, synthetic, nonbiodegradable polymers. PMMA has considerable versatility. It is used for dentures, arthroplasties, cranioplasties, and as cement for many orthopedic prostheses. PTFE has been used for augmentation (the Gore material SAM, subcutaneous augmentation material) and guided bone regeneration. The principle of guided bone regeneration is that new bone formation will occur by providing a "passageway" for osteoblast lineage cells and osteoblasts. Precluding soft tissue prolapse into an osseous deficit provides the underpinning to guide bone, thereby deterring scar formation. Physiologically, fibroblasts are more likely to populate an intraosseous deficit than osteoblasts; therefore, by sustaining a zone for migration of osteoblast lineage cells, the clinical outcome should be bone and not connective tissue scar. PMMA/PHEMA is commercially prepared under the name hard tissue replacement (HTR-MFI) as blocks and particulates. The block format is for augmentation whereas particulates have periodontal applications to restore deficient alveolar bone. However, no synthetic polymer alone can satisfy the conditions required for optimal scaffolding. Thus, new composite materials are being studied. Ceramics are also widely used in dental applications, and are being examined for bone tissue engineering applications. Two common ceramics used in dentistry and hip prostheses are alumina and hydroxyapatite. Alumina (Al2O3) has excellent corrosion resistance, good biocompatibility, high strength, and high wear resistance, and has been used for over 20 years in orthopedic surgery. Hydroxyapatite (HA) [Ca 10 (PO 4 ) 6(OH) 2] is a calcium phosphate based ceramic and has also been used for over 20 years in medicine and dentistry. HA is a major component of the inorganic compartment of bone. HA prepared commercially is biocompatible with biodegradability either absent or protracted. low degradation rate, can be controlled by varying the chemical structure major component of the inorganic compartment of bone a common bioceramic used for bone reconstitution 27

34 Chapter II. Bone and Non-Metallic Biomaterials exceptionally biocompatible with hard tissues highly osteoconductive and bioactive superb mechanical strength high osteoinductive potential possesses no cytotoxicity Fig. 2. Final fabricated HA scaffold with interconnected cylindrical pores. [31] Marshal Urist has defined the term osteoconduction as bone growth into and through a porous structure. Varying the chemical structure can control the degradation of hydroxyapatite. Tricalcium phosphate degrades much more quickly than HA. Medical grade tricalcium phosphate (TCP) was envisioned as an autograft expander. Plaster of Paris is the common name for the composition described chemically as calcium sulfate. The capacity to be fashioned and molded into the desired contour for the craniofacial complex is an asset. Unfortunately, unpredictable clinical outcome has been the track record for plaster of Paris. Finally, bioactive glasses have been shown to bind to soft tissue and bone. These bioactive glasses contain different ratios of Na2O-CaO-P2O5-SiO2. There are currently two commercially available glasses advertised for applications in bone sites. Native Polymers, or extracellular matrix proteins, are commonly exploited as bone graft materials. They offer good cytocompatibility and bioactivity. Collagens, which make up a majority of proteins in connective tissue such as bone, cartilage and tendons, are now widely used for biodegradable bone scaffolds. Various collagen-based products are currently under development. Since collagen is the most abundant protein in the extracellular matrix, it presents a chemical and physical environment favorable to bone recreation. Collagen is biocompatible, osteoconductive and biodegradable. Fig. 3. Collagen in connective tissue 28

35 Chapter II. Bone and Non-Metallic Biomaterials Collagen acts as a delivery system for bone morphogenetic proteins since it is: Fig. 4. Hierarchical Self-Assembly of Nano-Fibrils in Mineralized Collagen. [32] The organic phase of bone is principally type I collagen. When bone is demineralized with hydrochloric acid, the method used by most commercial venders, the bone derivative is largely type I collagen and a minimal per cent mixture of cell debris, a soup of soluble signaling molecules that are resistant to acidic demineralization, and residue ECM components. The format for the demineralized bone (DBM) can be either a range of particulate matter, blocks, or strips. Clinical reports on combinations of DBM and autograft favor this composition over DBM alone and underscore DBM as an autograft expander. Due to the strength limitations of collagen-based products like these, defect dimensions are limited and must be complemented by skeletal fixation. Gainfully, osteoinductive properties are imparted to these matrices when combined with bone marrow aspirates, thus greatly enhancing the repair process. The polysaccharide hyaluronic acid (Hy) appears to be a promising matrix material for bone tissue engineering, and is a glycosaminoglycan found in synovial fluid and cartilage. Hy has been experimentally determined to induce chondrogenesis and angiogenesis during remodeling and is being studied both individually and in combination with collagen as a matrix for bone repair. Finally, chondroitin sulfate is another glycosaminoglycan found in cartilage with potential applications as a bone tissue-engineered scaffold. Composites of ceramics and polymers are also widely studied. Composites can result in substitutes with properties between each of the respective materials. The greatest degree of success to obtain properties required for creating optimal scaffolding has been with biomaterials composed of collagen and HA. A biomaterial composed of HA powder and type I collagen, both from bovine origin demonstrates high osteoinductive and osteoconductive potential. For example, bovine collagen has been manufactured with HA. It is generally accepted that the combination of collagen and calcium-based ceramics provides a bone-like matrix that supports the adhesion, migration, growth, and differentiation of bone-forming cells. [2], [14]. 29

36 Chapter II. Bone and Non-Metallic Biomaterials Fig. 5.A biomaterial composed of HA powder and type I collagen Collagen has demonstrated to be the most effective component for it enhances adhesion and the guidance and contact guidance process of human osteoblasts positioning onto the collagen-ha composite surface. (a) Scanning electron microscopy image of bovine HA particles sintered at 1100ºC. Particles present irregular shapes. Bar: 500 microns. (b) Osteoblasts exhibiting cytoplasmatic projections strongly attached to HA surface. Bar: 500 microns.[1] II. 3. Modern non-metallic biomaterials for bone substitutes. Manufacture, properties, clinical applications and results. There are many commercial non-metallic biomaterials for bone substitutes available at present and their number is continuously growing. In this section will be detailed the properties, manufacture, clinical applications and results of some typical non-metallic biomaterials for bone substitutes used today. Typical examples: 1. Calcibon-synthetic bone substitute 2. Collapat II-Bone substitute biomaterial based on collagen and hydroxyapatite 3. Endobon 4. Palamed G the new generation bone cements 30

37 Chapter II. Bone and Non-Metallic Biomaterials II Calcibon-synthetic bone substitute Calcibon is a chemically derived synthetic bone substitute intended for the filling and reconstruction of bone defects. The Calcibon liquid is mixed with the Calcibon powder resulting in a smooth and malleable calcium phosphate paste, that has the ideal consistency for easy application in bone defects. This paste hardens at body temperature by an endothermic reaction to a microcrystalline, carbonated calcium-deficient hydroxyapatite. The chemical composition and the crystalline structure of the cured material mimics the chemical composition and crystalline structure of the mineral part of natural bone. Calcibon belongs to the family of calcium phosphates, a group of osteoconductive bone substitutes. Calcibon powder is synthesised from calcium and phosphate salts. The proportional distribution of the ingredients is almost equivalent to the calcium phosphate component of natural bone. a-tcp a-tri-calcium-phosphate CaHPO 4 calcium-hydrogen-phosphate CaCO 3 calcium-carbonate pha precipitated hydroxyapatite Calcibon liquid is an aqueous starting solution - di-sodium-hydrogen-phosphate solution -which initiates the curing process within the paste. The Indications: Calcibon is intended for refilling of bone voids caused by trauma or disease. Besides the augmentation the reconstruction of bone defects is a wide field of indications. Most common are: - distal radius fractures - tibial plateau fractures - calcaneus fractures. Calcibon is a one-time implant, which stays in place until it will be remodelled. It can be used in aseptic bone beds, except the area of open epiphyseal discs. Preconditions for use are an aseptic bone bed, accurate repositioning as well as proper reduction, stabilisation, and fixation of the defect. 31

38 Chapter II. Bone and Non-Metallic Biomaterials The physical and the biological features: 1. The Compressive Strength After hardening Calcibon will reach a very high compressive strength. In-vitro testing shows that after 6 hours the cured final material shows a compressive strength of approximately 25 MPa, already exceeding the strength of cancellous bone (10-20 MPa). This compressive strength continues to increase over time and results in a final strength of at least 60 MPa after 3 days. Thus the compressive strength is in the area of cortical bone ( MPa). 2. The Hardening Process After the complete filling of the bony defect with Calcibon the paste hardens in an endothermic reaction in about 10 minutes at body temperature. Thus there is no risk of heat necrosis development around the implant. 3. Compressive strength versus time. After 6 hours the compressive strength of the cured Calcibon is similar to that of the cancellous bone. After 3 days the final compressive strength of at least 60 MPa is reached. Calcibon hardens even in aqueous environment without falling apart. 4. The Biodegradability Calcibon has the potential to be biodegraded by the body and to be remodelled by the surrounding vital bone bed. The remodelling process is equilibrium of Calcibon break down by osteoclasts and the reconstruction of new bone by osteoblasts. Osteoclasts are specialised cells for degradation of existing bone or other potentially biodegradable material by extracellular enzymatic lysis. They release differentiating factors, which induce the differentiation of mesenchymal stem cells to activated osteoblasts. 32

39 Chapter II. Bone and Non-Metallic Biomaterials Osteoblasts are bone building cells synthesising and depositing new bone matrix with subsequent mineralization. Thereby the osteoblasts become enclosed in the new built bone as osteocytes. Osteocytes are fixed bone cells sensing the presence of bone damage and responsible for the constant repair of small lesions by self-repair mechanisms of the bone. The remodelling process depends on the individual and the environment where Calcibon is implanted. A vital bony bed, close contact with the surrounding bone, and a functional stimulation improves the rate of the remodelling. 5. The Biocompatibility In-vitro tests showed that Calcibon has no mutagenic, no cytotoxic, and no genotoxic potential. After in-vivo implantation, Calcibon showed neither acute nor sub-acute systemic toxicity, no sensitising potential was observed nor the material caused neither sub-cutaneous nor intra-cutaneous irritation. Osteoclasts recognise bone mineral like surface Osteoclasts start degradation Osteclasts release differentiating factors and thereby induce differentiation (a) and activation (b) of osteoblasts Osteoblasts synthesise and deposit new bone matrix with subsequent mineralization Calcibon-animal studies: Trabecular Bone Response The aim: investigation of the physicochemical, biological, and handling properties of Calcibon when implanted in trabecular bone. 33

40 Chapter II. Bone and Non-Metallic Biomaterials The design: Calcibon was implanted as a paste into the femoral trabecular bone of goats for 3 days, 2, 8, 16, and 24 weeks, respectively. Polymethylmethacrylate implants served as control. The results: -Calcibon was easy to handle and fast setting with good cohesion when in contact with body fluids. -Histological evaluation after 2 weeks showed abundant bone apposition on the Calcibon surface without any inflammatory reaction or fibrous encapsulation. -At later time points, the Calcibon implants were totally covered by a thin layer of bone. Osteoclast-like cells were clearly degrading the bone substitute. -At all time periods the PMMA cement was surrounded by a fibrous membrane without any evidence of PMMA-bone contact. Fig. 1. Detail of bone formation at the interface after 2 weeks. Ob, osteoblasts; Oc, osteocyte; *, osteoid layer. Bar = 125 urn. The Calcibon surface is already almost completely covered with a thin layer of woven bone. This newly formed bone is characterised by the presence of a layer of osteoid lined with osteoblasts. 34

41 Chapter II. Bone and Non-Metallic Biomaterials (a) (b) (a) After 24 weeks of implantation in trabecular bone. Bar = 250 μm. Continued biodegradation of Calcibon and subsequent bone ingrowth. The bone is very mature and cannot be discerned from original trabecular bone. (b) Low magnification photomicrograph of a transversal section of Calcibon implant after 24 weeks of implantation. Bar = 1500 μm. Parts of the bone substitute mass have been biodegraded followed by bone ingrowth. Remodelling activity at the interface, as characterised by the presence of remodelling lacunae and osteoclast-like cells, was still seen. Measurements of the implant area suggest a decrease in Calcibon area at 16 and 24 weeks of implantation, though this was not statistically significant. Cortical Bone Response The aim: investigation of the physicochemical and biological properties of Calcibon when implanted in cortical bone. The design: Calcibon was injected as a paste into tibial cortical bone defects in goats. The animals were sacrificed after 3 days, 2, 8, 16, and 24 weeks, respectively. Polymethylmethacrylate implants were used as control. The results: - Light microscopy evaluation showed that after 2 weeks the Calcibon was in tight contact with the bone without any inflammatory reaction or fibrous encapsulation. - At later time points, the Calcibon implants were totally covered by a thin layer of bone, and osteoclasts were clearly degrading the bone substitute and depositing new bone. - Although the bulk of the material was still in situ after 24 weeks, the progressive osteoclast degradation of Calcibon followed by new bone formation suggests that all of the material might be replaced. - In contrast to Calcibon, a thin fibrous capsule surrounded all the PMMA implants. The animal studies show that Calcibon is biocompatible, osteoconductive as well as osteotransductive, i.e. biodegradation of the material takes place, directly followed by the formation of new bone. The Clinical Study: The aim: evaluation of the safety and performance of Calcibon in filling cancellous bone defects as well as the long time outcome. Safety - frequency of adverse events was evaluated. Performance - osseous integration, operative handling, and clinical results were monitored. Long time outcome - X-rays have been evaluated for osseous integration and loss of position of the implant, after 1 year. The design: A multi-centre, prospective, open, uncontrolled, non-randomised clinical study has been performed in 3 centres. An overall number of 44 patients with cancellous bone defects of any indication, who need an operative filling, have been treated. Only closed fractures were treated where in the traditional operation method the surgeon would have implanted autologous spongiosa. The follow up period was 1 year. 35

42 Chapter II. Bone and Non-Metallic Biomaterials The results: - 44 patients were included in the study, 25 females and 19 males with a mean age of 50 years (16-82 y). - Different quantities of Calcibon were used, from 2 g up to 30 g (mean 8.1 g). - In 41 cases trauma defects were treated: - 19 distal radius fractures, - 11 tibial head plateau fractures, - 7 calcaneus fractures, - 1 humerus fracture, - 1 lower leg fracture, - 1 phalanx fracture, and - 1 pilon fracture. The clinical results 1 year after implantation of Calcibon were comparable to the results known from autograft. There were no disturbances in bone healing, no radiolucent zones around the Calcibon implant on X-rays, neither dislocations nor fractures of the Calcibon implants, and no problems during handling and application of Calcibon. The clinical study shows that: - 91 % of the surgeons found the treatment with Calcibon comparable or better to the conventional treatment, - 97 % of the cases Calcibon showed a good or excellent performance, -100 % of the cases the mixing and application of Calcibon was easy even in the presence of blood, and -100 % of the cases Calcibon could be well adapted to the defect. The Treated Fracture Types: 44 patients showed the following distribution of fracture types in percent (%): 36

43 Chapter II. Bone and Non-Metallic Biomaterials Calcibon-the mixing procedure and the presentations: The processing cycle has been optimised in consideration of the operating Room environment: Short mixing time 1 minute mixing phase Sufficient application time 4 minutes processing phase Acceptable final setting time 5 minutes curing phase Overall required time Σ 10 minutes After 1 minute of thorough mixing, there are 4 minutes for processing and applying Calcibon. During the curing phase every further manipulation must be avoided to allow a proper crystalline organisation of the material.. The mixing Instructions: Mixing instructions on the sterile blister for the use in the operating room. The paste is very smooth and malleable, easy to applicate and mix, having excellent handling properties. [15] 37

44 Chapter II. Bone and Non-Metallic Biomaterials II Collapat II- Bone substitute biomaterial based on collagen and hydroxyapatite COLLAPAT II is a haemostatic bone substitute material presented as a fleece. It is composed of a collagen structure in which ceramised hydroxyapatite granules are dispersed. Description, Indications, Advantages: Description: The hydroxyapatite granules give Collapat II its osteoconductive properties. They are biodegradable. The collagen gives Collapat II its strong haemostatic power. It is completely resorbable in a few weeks. It is extracted from calf skins. The manufacturing procedure comprises stages recognised to inactivate viruses and non-conventional transmissible agents. These treatments make it possible to ensure maximum microbiological safety for Collapat II, in particular with respect to the agent responsible for BSE. The collagen incorporated into Collapat II also enters into the composition of numerous medical devices approved in Europe, Canada and the United States. Indications: In orthopaedics: Collapat II is used to promote the repair of various types of bone lesions: lifter removal of corticocancellous bone fragments. Biter tumour resection. It revision cases. B surgical spondylodeses. B cases of pseudarthrosis. Collapat II is also used to support bone formation in maxillofacial surgery and stomatology. Advantages: - is composed of two of the main constituents of the bone, collagen and hydroxyapatite. - is osteoconductive- it enables rapid and multicentric bone regeneration in the collagen network by means of the finely dispersed and distributed small hydroxyapatite granules. - is haemostatic- when applied in haemorrhagic lesions, stops the blood flow in a few minutes. - is biodegradable-due to the tiny particle size the hydroxyapatite granules are resorbed. - is malleable-due to its spongy structure, the product takes the consistency of a paste in contact with blood or tissue fluids, enabling optimal adaptation to the shape of the bone surface. - can be impregnated with an antibiotic solution in the event of septic bone infection. - makes it possible to avoid costly autologous or allogenic transplantation operations. 38

45 Chapter II. Bone and Non-Metallic Biomaterials Performance study: The osteoconductive properties of Collapat II were demonstrated by means of a study in rabbits. Was implanted in a corticocancellous defect 4.2 mm in diameter in the femoral site. The progression of bone regrowth was evaluated up to 3 months by histological sections from specimens included in resin magnified by a factor of 20. In the next image is presented: TO: Implantation date The hydroxyapatite granules are dispersed in the collagen matrix. T1: Primary bone. After 1 month, the collagen has been completely resorbed and the implanted sites are largely colonised with primary neo-formed bone tissue, with signs of direct bone hydroxyapatite apposition. T2: Mature bone After 3 months, the bone reconstruction is complete and is characterized by mature remodelled bone tissue. Package: Collapat II is presented packaging, impermeable to light, in unit box form, sterilised by β-rays. [16] 39

46 Chapter II. Bone and Non-Metallic Biomaterials II Endobon Manufacture, Properties, Application: Endobon is a natural hydroxyapatite ceramic (HA-ceramic), which is particularly suitable for use as a bone substitute material. It is of biological origin and osteoconductive. The inner structure with its interconnecting system of macro and micro pores with a pore size of 100 to 1,500 μm allows the newly formed bone to grow directly on to the ceramic surface through the whole implant. This leads finally to a stable osseous integration of the Endobon implant, no matter whether blocks, cylinders, or granules are implanted. Is manufactured in a two-stage high-temperature process: - Pyrolysis at a temperature above 900 C, - Sintering at a temperature above C. This leads to a combustion of all organic material in the bone, thus ensuring complete de-proteinisation and hence destruction and elimination of all bacteria, viruses from the original material. 40

47 Chapter II. Bone and Non-Metallic Biomaterials Clinical use: Hydroxyapatite ceramics are in general an excellent alternative treatment to autologous cancellous bone graft. Especially for traumatic fractures (tibial head fractures, fractures of the calcaneus, and of the distal radius) reduces the operation time and the anaesthesia time. The potential risks of taking autologous cancellous bone graft can be avoided. Beside traumatic bone defects even bone cysts and non-infected bone cavities of any other genesis can be filled. Patient: 55-year old Diagnosis: Joint depression fracture of calcaneous a. pre-operative image b. operative treatment with an H-plate in combination with osteosynthesis and filling with Endobon c. 10 months post-operatively, after metal implant removal d. sufficient function of the right talo-calcaneal joint (3 and 10 months postoperatively, respectively) 41

48 Chapter II. Bone and Non-Metallic Biomaterials Restrospective analysis of fracture healing: Endobon was used to fill metaphyseal defects associated with proximal tibia fractures. After reconstruction of the joint surface the defect was filled completely with Endobon blocks or cylinders and stabilised by internal plate fixation subsequently. Patients: n=35 Diagnosis: Fractures of proximal tibia Follow up: up to 91 months X-ray pictures of a 64-year old female patient with an uni-condylar proximal tibia fracture (AO type 41 B3) after defect filling with Endobon (a.-p. and lateral). a. pre-operative images b. day 1 post operatively c. 6 months post-operatively d. 57 months post-operatively Even after removal of the metal implants no secondary loss of bony reconstruction could be seen. Results after proximal tibia fractures: X-ray images of a 75-year old female patient with a bicondylar proximal tibia fracture (AO type 41 C3) after defect filling with 3 Endobon blocks (a. -p. and lateral). a. pre-operative images b. day 1 post-operatively c. 6 months post-operatively d. 62 months post-operatively 42

49 Chapter II. Bone and Non-Metallic Biomaterials The fracture is completely consolidated by new bone with good functional results. At the margins of the Endobon ceramic an increasing radiolucency could be observed. Function, radiology and histology: Function: A secondary loss of reduction due to mechanical failure of the Endobon implant did not occur even after removal of the metal implants. Radiology: Bony consolidation of the fracture occurred in the usual time course. During the healing process, the boundary zones of the implanted Endobon developed a progressive radiolucency indicating an increasing osseous integration. Histology: The newly formed bone grows directly on to the porous structure of Endobon without a fibrous tissue interface. Even within the pores bone trabeculae and an accompanying vascularization could be observed. [17] 43

50 Chapter II. Bone and Non-Metallic Biomaterials II Palamed G the new generation bone cements Palamed and Palamed G have been developed from Palacos R to meet the needs of Modern Cementing Technique. Palamed and Palamed G are high-viscosity cements that have an initially reduced viscosity, permitting them to be used, without pre-chilling, in modern vacuum mixing systems. This standardized procedure, in addition to saving valuable time in the operating room, produces homogeneous cement with reduced porosity. Antibiotic-loaded cements reduce the risk of infection in both primary and revision operations. Palamed G is frequently used for primary THR and TKR because of the long-term savings it affords, as compared to the results of using plain bone cement, with its increased risk of revision. Provides excellent antibiotic release. Palamed and Palamed G, contain chlorophyll, which gives the cement its characteristic green colour used for easy distinction from bone and body tissue during surgery. Palamed G comes in three pack sizes: 20 g, 40 g and 60 g. Palamed comes in size 40 g. Modern bone cements must meet a multitude of international technical and medical standards. Scandinavian publications show that high viscosity bone cements yield better longterm results than low viscosity cements. Although low viscosity cements are easier to mix, the difficulties associated with their application may be the reason for their inferior performance. Palamed advantages: - Based on Palacos R - Excellent mechanical properties - Long-lasting local gentamicin release - Quality control of every batch - Optimal handling in modem vacuum mixing systems - Pre-chilling unnecessary - Green colour for easy recognition during surgery - Reduced amount of zirconium - Part of the complete Modern Cementing Technique concept ISO Standard: Following the implantation of a cemented endoprosthesis, bone cement is subjected to high mechanical stress. The mechanical properties of bone cement should therefore be tested for compressive strength, bending strength, and Young's modulus according to ISO Both Palamed and Palamed G exceed the international standards established for strength. As the cemented implant is subjected to dynamically alternating loads, long-term stability of the cement is critical. Because of the cyclic stresses bone cement is subjected to in vivo, fatigue properties of bone cement are an important factor in the long-term survival of a cemented hip replacement. 44

51 Chapter II. Bone and Non-Metallic Biomaterials Antibiotic-loaded cement for decreased risk of revision: The addition of antibiotics to bone cement was first undertaken by Buchholz. Almost 75% of all bacteria that can be isolated during hip operations are gram-positive, with staphylococci having the greatest share. Among gram-negative bacteria, E. coli and pseudomonas are most common. Gentamicin has proven to be the antibiotic of choice for bone cement, as its broad therapeutic spectrum covers gram-positive and gram-negative bacteria. Is bactericidal on proliferating and resting pathogens, and its release from the bone cement is superior to that of other antibiotics. A protracted release of gentamicin from cured bone cement was proven as early as Studies have shown that Palamed G provides high local concentrations of gentamicin over several days. In recent years a considerable number of case studies have been published confirming the successful use of antibiotic-loaded cements. Antibiotic-loaded cements such as Palamed G provide effective gentamicin concentrations in the tissue surrounding the implant. 45

52 Chapter II. Bone and Non-Metallic Biomaterials Additional clinical advantages of gentamicin-loaded cements include low systemic and urinary antibiotic concentrations and a consequent lack of ototoxic and nephrotoxic complications. Staphylococcus coag.negative Enterobacteriaceae Staphylococcus aureus Staphylococcus sp. Anaerobics Pseudomonas sp Others For mixing and delivery of Palamed and Palamed G, it is recommended to use a closed vacuum mixing system, which helps achieve reproducible, homogenous bone cement of the highest quality. Handling properties are highly dependent on the temperature of the bone cement and of the operating room. Higher temperatures make for a shorter working phase and a faster setting time. Examples of handling schedules for different mixing methods and temperatures. 46

53 Chapter II. Bone and Non-Metallic Biomaterials General information: Before using Palamed and Palamed G, the surgeon should be thoroughly familiar with the properties, handling and use of these cements and should have read all relevant literature. After preparation of the bone bed or immediately after introduction of the cement and prosthesis, pressure increase in the medullary canal may cause a temporary fall in blood pressure. In addition to hypotension, pulmonary embolism and cardiac arrest have been encountered in rare cases. These cardiovascular and respiratory side effects, known as bone implant syndrome, are mainly caused by rise of intramedually pressure during preparation of the implant bed and application of cement and implant. The site of the prosthesis should be rinsed thoroughly with an isotonic solution (e.g. physiological saline) before implantation To minimize the rise in pressure in the medullary canal during implantation of the cement and the prosthesis, sufficient drainage is recommended. In the presence of pulmocardiovascular disturbance, the blood loss should be monitored and aneasthesiological measures may be required (e.g. in the event of acute respiratory failure). Palamed G and Palamed must not be used in cases of known hypersensitivity to gentamicin or to other constituents of the bone cement. [18] II. 4. Bone cement and modern cementing technique. The elements of Modern Cementing Technique: Bone cement Mixing and delivery Bone bed preparation Pressurisation Revision Safe working environment and potential hazards Evolution: Thanks to the research and groundbreaking thesis of Otto Rohm, PMMA cements came into use for dental applications in In 1943 Kulzer patented a cold curing process for PMMA, which led to the production of PMMA bone cements in the 1960s. In the early 1960s Sir John Charnley introduced the concept of low friction arthroplasty. His system of components, with a metal implant for the femur and a plastic device for the acetabulum - both fixed with PMMA bone cement - is still commonly used for total hip replacement. In the late 1960s Buchholz investigated with the German companies Kulzer and E. Merck the addition of antibiotics to bone cements. Numerous clinical studies at the Endo- Klinik in Hamburg followed, leading to the first antibiotic-loaded bone cement, Refobacin-Palacos R. In the 1980s new techniques, aimed at improving the properties of bone cement, were introduced. Vibration, centrifugation and mechanical mixing were studied, but none of these methods gained widespread acceptance. Then mixing under vacuum was introduced, which after some refinement made it possible to produce high quality, homogeneous bone cement. Prof Lars Lidgren of Lund University Hospital was a pioneer 47

54 Chapter II. Bone and Non-Metallic Biomaterials and a teacher in the field of closed mixing and new generation cementing techniques using high viscosity cement, and one of the first vacuum mixing systems was developed by the Swedish firm MITAB, now Biomet Cementing Technologies AB. In 1981 Joe Miller introduced the concept of micro-interlock, which is characterised by a stronger bone cement interface. Several factors are responsible for the improved mechanical quality of micro-interlock: the bone bed is carefully cleaned, a high and reproducible bone cement quality is achieved by mixing and collecting under vacuum and pressurisation devices are used for cement filling. In the past, cementing technique comprised hand mixing and finger packing. Today's modern technique benefits from vacuum mixing and the use of distal femoral plugs, femoral and acetabular pressurisation, bone bed brushing and high-pressure pulse lavage. Each step of Modern Cementing Technique has been linked to an approximate 20% reduction in revisions for aseptic loosening. Clean bone bed The use of brushes and high-pressure pulse lavage improves bone cement interface and prevents haemodynamic circulatory changes. Safe working environment A closed mixing and : delivery system reduces monomer fumes and minimises skin contact. Good cement filling Pressurisation and distal plugs improve bone cement interface and yield good stress distribution. High and reproducible bone cement quality Mixing and collection of bone cement under vacuum reduces porosity and improves fatigue strength 48

55 Chapter II. Bone and Non-Metallic Biomaterials Antibiotic-loaded A broad spectrum of high and low viscosity bone cements (with or without antibiotics) ensures optimal clinical results by enabling the proper cement to be matched with the necessary cementing technique. High viscosity bone cements have been linked to a lower risk of revision and aseptic loosening in total hip arthroplasties and are the gold standard for treating total hip replacement (THR) and total knee replacement (TKR). Low viscosity bone cements are increasingly being used for special applications. High viscosity PLAIN Low viscosity Antibiotic-loaded cements reduce the risk of infection in both primary and revision operations. These cements are frequently used for primary THR and TKR because of the longterm savings they afford as compared to the results of using plain bone cement, with its increased risk of revision. 49

56 Chapter II. Bone and Non-Metallic Biomaterials Gentamicin and clindamycin All Biomet antibiotic-loaded cements contain gentamicin. For one- and two-stage revisions, revisions for infection and procedures on risk patients, a double-loaded antibiotic bone cement, such as Copal (with gentamicin and clindamycin), is recommended. For modern mixing systems, specially developed cements such as Palarned further improve the handling technique in cemented arthroplasties. Biomet's standard bone cements cover most cases when cemented arthroplasty is necessary. Sometimes, however, standard cements are not adequate. In these cases, Biomet can offer a patient matched bone cement manufactured according to European medical device directive specially made for a certain patient. Biomet cements contain chlorophyll, with a characteristic green colour for easy distinction from bone and body tissue during surgery. Bone bed preparation-the vital preparatory stage: Before and after cleaning with high-pressure pulse lavage Increased cement penetration for highpressure pulse lavage compared to syringe lavage. -Increased contact area between bone and cement -A clean and dry bone cavity -High-pressure pulse lavage for efficient cleaning 50

57 Chapter II. Bone and Non-Metallic Biomaterials In the early days of cemented total hip arthroplasty (THA}, preparation of the cancellous bone stock was given relatively little attention. To achieve adequate cement interdigitation, anchoring holes in the acetabulum and sharp cutting broaches in the femur, which help to preclude bone compaction, have been shown to be beneficial. Particularly, preservation and subsequent meticulous cleansing of the cancellous framework are of utmost importance. A clean, dry bone cavity improves the quality of the bone cement interlock. Accidental entrapment of blood, bone marrow and tissue debris in the cement may create defects, compromising both cement and interface strength. Brushes may be used to remove soft tissue and loose cancellous bone. Pulsed lavage is commonly used in trauma cases to remove debris, blood and dead tissue. Irrigation with high pressure pulsed lavage causes no damage to healthy tissue and makes for very efficient cleaning. Its use should also be considered mandatory in cemented joint arthroplasty to clean the spongious bone. It has been shown that the rise of intramedullary pressure caused by the insertion of cement and prosthesis can lead to fat and bone marrow intravasation. This phenomenon may provoke a circulatory reaction, with a decrease in peripheral blood pressure and impaired central circulatory and pulmonary functions. Thorough pulsed lavage, including the removal of medullary contents prior to retrograde cement application, is a proven important and logical step in reducing the risk of fat embolism. To perform a successful and long-lasting cemented THA, the surgeon needs to be prepared to strictly implement modern techniques of bone preparation, the use of pulsatile lavage and pressurised application of cement. Choosing the right vacuum mixing system-palacos R mixed by hand, in a common mixing system, and in the Optivac. Modem closed vacuum mixing systems Reduced monomer exposure, standardised mixing in the delivery system. Cycles to failure Hand mixing Vacuum mixing 51

58 Chapter II. Bone and Non-Metallic Biomaterials The presence of large pores in the cement may lead to a rapid propagation of cements cracks. The fatigue life improves ten times by vacuum mixing of Palacos (at 4 C) rather than hand mixing. Mixing and delivery: The development of cement mixing and delivery systems has been governed by three key objectives: standardising the cementing technique, improving cement quality and minimising monomer exposure. A hip joint is subjected to forces as great as ten times body weight and is set in motion up to two million times each year. So the fatigue properties of cement are critical. Porosity has been found to decrease the cement's mechanical strength. Pores or other inclusions concentrate stress in the material and initiate fatigue cracks, which may lead to failure. Other factors contributing to poor bone cement behaviour are molecular weight, cement viscosity, type of cement and implant. While vacuum mixing reduces microporosity, reports show that fatigue fractures occur at the largest macropores. Macroporosity is greatly effected by choice of vacuum mixing system. Studies have shown that the Optivac system, with its patented design for vacuum collection, produces a cement of much lower macroporosity than cement mixed in other vacuum mixing systems. A non-homogeneous cement mix with unbonded particles may also compromise cement quality. Hand mixing with a spatula and bowl, followed by transferral to a delivery system, makes it difficult to achieve a homogeneous cement mix. Modern mixing and delivery systems produce reproducible, homogeneous cement of consistently high quality. A system with a cement delivery gun produces better pressurisation of cement into bone bed than does a system with manual delivery. Closed, pre-packed systems, which offer enhanced ease-of-use and reduced exposure, are the future of cement mixing and delivery. Systems pre-filled with polymer powder are the first step in this direction. Closed vacuum mixing system Optivac Pressurization Pressurisation has been shown to afford greater penetration into cancellous bone, reduce bone cement porosity, improve the interlock of the bone cement interface and enhance cement strength. 52

59 Chapter II. Bone and Non-Metallic Biomaterials For good intrusion of the bone cement into the trabeculae, cement viscosity is critical. If viscosity is too low, blood pressure may force the cement out of the bone cavity and cause blood laminations in the cement. Ideal viscosity is high enough to prevent the cement from mixing with blood or fat/bony material yet low enough to penetrate the bone adequately. Retrograde filling High-pressure to achieve microinterlock Stem in neutral position Improved cement penetration Cement application Acetabular pressurisation Even cement mantle Cement mixing in a bowl followed by finger packing does not provide adequate fill, especially in the femoral canal. Retrograde cement filling in the femur reduces risk of revision. This technique prevents air voids during cement application, reducing porosity. Lamination is also reduced, making it possible to achieve a completely cement-filled cavity. Pressurisers are available for acetabular, femoral and knee pressurisation. They should be used to maintain pressure on the bone cement until it is doughy enough to resist the force of blood pressure. In order to achieve good filling and pressurisation in hip and knee arthroplasties, a cement restrictor may be used to plug the shaft. Pressurisation and lavage of the cancellous bone significantly improves cement penetration. The bone cavity should be shaped to provide an even cement mantle between the bone and the prosthesis. The cement mantle should be approximately 2-3 mm thick. 53

60 Chapter II. Bone and Non-Metallic Biomaterials Normally 40 g of cement is used for acetabulum and g for femur. A stem centralizer can help guide the femoral prosthesis to a neutral position in the cement. An even cement mantle means better stress distribution and reduced risk of cement mantle failure. Hip pressurisation Knee pressurisation Cement restrictors Copal cement contains gentamicin and clindamycin, a combination known to have a bactericidal effect on more than 90% of the bacteria common to infected arthroplasty cases. The advantage of antibiotic-loaded cements is their local release of high antibiotic concentrations at the site of implantation. The local release of antibiotics causes low rise in serum or urine concentrations and therefore does not produce the side effects associated with systemic antibiotics. The high initial and subsequent protracted antibiotic release of the bone cement provides effective protection of the implant. In many countries antibiotic-loaded bone cements are used in primary operations because of the reduced risk of revision and consequent long-term savings. Double-loaded antibiotic bone cement provides extra security in cases such as one- and two-stage revisions, revisions for infections and risk patients with severely decreased immunity. One- and two-stage revision: Revision procedures are performed in one or two sessions. During one-stage hip revision surgery, the implant and the former cement mantle is removed. In a two-stage revision, a temporary antibiotic cement joint spacer may be used for local antibiotic protection. This temporary cement spacer prevents contraction of soft tissue and stabilises the infected joint. 54

61 Chapter II. Bone and Non-Metallic Biomaterials Antibiotic-loaded bone cement Two-stage revision Cement removal instruments (a) (b) (a) Closed system minimizes monomer fumes and skin contact. Modern vacuum mixing systems remain closed at every step, from adding the cement components to cement application (b) One cartridge for mixing and delivery No need to transfer bone cement chamber to delivery cartridge. Minimize skin contact An extra pair of PE gloves gives the best available protection against monomer exposure 55

62 Chapter II. Bone and Non-Metallic Biomaterials Safe working environment Increased awareness of the adverse effects of MMA, combined with growing demand for safety in the workplace, have led to the development of closed mixing systems, which minimise exposure to monomer fumes in the OR and prevent direct contact with the cement. Liquid monomer is a volatile and flammable solvent and should not come into direct contact with the skin. Exposure to MMA may cause irritation of the upper respiratory tract and the eyes. Indications are headache, nausea and lack of appetite. A modern, closed vacuum mixing system should be used to prevent direct contact, and special gloves should be worn during cementation. Law in many countries regulates monomer exposure levels. In the UK, for example, regulations allow for a maximum daily MMA exposure of 208 mg/m 3 and a maximum short-term exposure of 416 mg/m 3. Monomer odours may be sensed even when the levels are exceedingly low. Closed systems such as the Optivac reduce MMA exposure to barely detectable levels - far below the safe limits established by law. Hand-mixed cement contains a significant number of pores, which reduce strength and compromise implant survival. Bone cement mixed and collected under vacuum has proven to be practically non-porous, resulting in greater durability. The need for revision surgery with is associated costs is greatly reduced. A closed vacuum mixing system reduces exposure to monomer fumes and eliminates direct contact with bone cement during mixing and delivery. The working environment for the operating staff is improved, and the risk of fume-induced headaches, respiratory irritation and allergic reactions becomes almost non-existent. An easy-to-use modem mixing and delivery system achieves reproducible, highquality bone cement at all times. [19] 56

63 Chapter III. Bioresorbable Biomaterials in Orthopaedics Chapter III Bioresorbable Biomaterials in Orthopaedics General Most of the biological components constructing our body are repeatedly undergoing bioabsorption and biosynthesis in our life cycle. A typical example is remodeling of bones. This is in marked contrast with most biomaterials that will stay permanently in our body without biodegradation and resorption. Resorbable biomaterials, on the other hand, gradually disappear from the body as a result of hydrolysis. Historically, reconstructive implants have been made of permanent materials such as metal or ceramic. Permanent or nonresorbable materials remain in the body after healing takes place, unless they are surgically removed. In an effort to provide patients alternative treatment options, surgeons and scientists have explored the benefits of bioresorbable materials made from molecules similar to those in the human body, which resorb while the tissue is healing. This chapter is a review of the chemistry of the polymers, including synthesis and degradation, describe how properties can be controlled by proper synthetic controls such as copolymer composition, highlight special requirements for processing and handling, and presents some of the commercial devices based on these materials. In the first half of this century, research into materials synthesized from glycolic acid and other -hydroxy acids was abandoned for further development because the resulting polymers were too unstable for long-term industrial uses. However, this very instability leading to biodegradation has proven to be immensely important in medical applications over the last three decades. Polymers prepared from glycolic acid and lactic acid have found a multitude of uses in the medical industry, beginning with the biodegradable sutures first approved in the 1960s. Since that time, diverse products based on lactic and glycolic acid and on other materials, including poly(dioxanone), poly(trimethylene carbonate) copolymers, and poly (-caprolactone) homopolymers and copolymers have been accepted for use as medical devices. In addition to these approved devices, a great deal of research continues on polyanhydrides, polyorthoesters, polyphosphazenes, and other biodegradable polymers. There may be a variety of reasons for which a medical practitioner wants a materials to degrade, but the most basic begins with the physician's simple desire to have a device that can be used as an implant and will not require a second surgical intervention for removal. Besides eliminating the need for a second surgery, the biodegradation may offer other advantages. For example, a fractured bone that has been fixated with a rigid, nonbiodegradable stainless implant has a tendency for refracture upon removal of the implant. Because the stress is borne by the rigid stainless steel, the bone has not been able to carry sufficient load during the healing process. However, an implant prepared from biodegradable polymer can be engineered to degrade at a rate that will slowly transfer load to the healing bone. 57

64 Chapter III. Bioresorbable Biomaterials in Orthopaedics Polymer chemistry Biodegradable polymers can be either natural or synthetic. In general, synthetic polymers offer greater advantages than natural materials in that they can be tailored to give a wider range of properties and more predictable lot-to-lot uniformity than can materials from natural sources. Synthetic polymers also represent a more reliable source of raw materials, one free from concerns of immunogenicity. Table I. Properties of common biodegradable polymers. Polymer Melting Point ( C) Glass-Transition Temp ( C) Modulus (GPa) Degradation Time (months) PGA to 12 LPLA >24 DLPLA Amorphous to 16 PCL ( 65) ( 60) 0.4 >24 PDO N/A ( 10) to 12 PGA-TMC N/A N/A to 12 85/15 DLPLG Amorphous to 6 75/25 DLPLG Amorphous to 5 65/35 DLPLG Amorphous to 4 50/50 DLPLG Amorphous to 2 a. Tensile or flexural modulus. b. Time to complete mass loss. Rate also depends on part geometry. The general criteria for selecting a polymer for use as a biomaterial is to match the mechanical properties and the time of degradation to the needs of the application (see Table I). 58

65 Chapter III. Bioresorbable Biomaterials in Orthopaedics The ideal polymer for a particular application would be configured so that it: Has mechanical properties that match the application, remaining sufficiently strong until the surrounding tissue has healed. Does not invoke an inflammatory or toxic response. Is metabolised in the body after fulfilling its purpose, leaving no trace. Is easily processable into the final product form. Demonstrates acceptable shelf life. Is easily sterilized. The factors affecting the mechanical performance of biodegradable polymers are those that are well known to the polymer scientist, and include monomer selection, initiator selection, process conditions, and the presence of additives. These factors in turn influence the polymer's hydrophilicity, crystallinity, melt and glass-transition temperatures, molecular weight, molecular-weight distribution, end groups, sequence distribution (random versus blocky), and presence of residual monomer or additives. In addition, the polymer scientist working with biodegradable materials must evaluate each of these variables for its effect on biodegradation. Biodegradation has been accomplished by synthesizing polymers that have hydrolytically unstable linkages in the backbone. The most common chemical functional groups with this characteristic are esters, anhydrides, orthoesters, and amides. The following section presents an overview of the synthetic biodegradable polymers that are currently being used or investigated for use in wound closure (sutures, staples); orthopedic fixation devices (pins, rods, screws, tacks, ligaments); dental applications (guided tissue regeneration); cardiovascular applications (stents, grafts); and intestinal applications (anastomosis rings). Most of the commercially available biodegradable devices are polyesters composed of homopolymers or copolymers of glycolide and lactide. There are also devices made from copolymers of trimethylene carbonate and -caprolactone, and a suture product made from polydioxanone. Polyglycolide (PGA). Polyglycolide is the simplest linear aliphatic polyester. PGA was used to develop the first totally synthetic absorbable suture, marketed as Dexon in the 1960s by Davis and Geck. Glycolide monomer is synthesized from the dimerization of glycolic acid. Ring-opening polymerization yields high-molecular-weight materials, with approximately 1 3% residual monomer present. PGA is highly crystalline (45 55%), with a high melting point ( C) and a glass-transition temperature of C. Because of its high degree of crystallization, it is not soluble in most organic solvents; the exceptions are highly fluorinated organics such as hexafluoroisopropanol. Fibers from PGA exhibit high strength and modulus and are too stiff to be used as sutures except in the form of braided material. Sutures of PGA lose about 50% of their strength after 2 weeks and 100% at 4 weeks, and are completely absorbed in 4 6 months. Glycolide has been copolymerized with other monomers to reduce the stiffness of the resulting fibers. 59

66 Chapter III. Bioresorbable Biomaterials in Orthopaedics Polylactide (PLA). Lactide is the cyclic dimer of lactic acid that exists as two optical isomers, d and l. l-lactide is the naturally occurring isomer, and dl-lactide is the synthetic blend of d-lactide and l-lactide. The homopolymer of l-lactide (LPLA) is a semicrystalline polymer. These types of materials exhibit high tensile strength and low elongation, and consequently have a high modulus that makes them more suitable for loadbearing applications such as in orthopedic fixation and sutures. Poly(dl-lactide) (DLPLA) is an amorphous polymer exhibiting a random distribution of both isomeric forms of lactic acid, and accordingly is unable to arrange into an organized crystalline structure. This material has lower tensile strength, higher elongation, and a much more rapid degradation time, making it more attractive as a drug delivery system. Poly(l-lactide) is about 37% crystalline, with a melting point of C and a glass-transition temperature of C. The degradation time of LPLA is much slower than that of DLPLA, requiring more than 2 years to be completely absorbed. Copolymers of l-lactide and dl-lactide have been prepared to disrupt the crystallinity of l-lactide and accelerate the degradation process. PLA or polylactide, is prepared from the cyclic diester of lactic acid (lactide) by ring opening polymerization as shown below: Poly(-caprolactone). The ring-opening polymerization of -caprolactone yields a semicrystalline polymer with a melting point of C and a glass-transition temperature of 60 C. The polymer has been regarded as tissue compatible and used as a biodegradable suture in Europe. Because the homopolymer has a degradation time on the order of 2 years, copolymers have been synthesized to accelerate the rate of bioabsorption. For example, copolymers of -caprolactone with dl-lactide have yielded materials with more-rapid degradation rates. Ethicon is selling a block copolymer of -caprolactone with glycolide, offering reduced stiffness compared with pure PGA, as a monofilament suture, under the trade name Monacryl. Polycaprolactone is synthesized from e-caprolactone as shown below: 60

67 Chapter III. Bioresorbable Biomaterials in Orthopaedics This semi-crystalline polymer absorbed very slowly in vivo and released e- hydroxycaproic acid as the sole metabolite. Nonenzymatic bulk hydrolysis of ester linkages followed by fragmentation and release of oligomeric species. Fragments ultimately scavenged by macrophages and giant cells. Amorphous regions of the polymer are degraded prior to breakdown of the crystalline regions. Copolymers of e-caprolactone and L-lactide are elastomeric when prepared from 25% e-caprolactone, 75% L-lactide and rigid when prepared from 10% e-caprolactone, 90% L-lactide. Poly(dioxanone) (a polyether-ester). The ring-opening polymerization of p- dioxanone resulted in the first clinically tested monofilament synthetic suture, known as PDS (marketed by Ethicon). This material has approximately 55% crystallinity, with a glass-transition temperature of 10 to 0 C. The polymer should be processed at the lowest possible temperature to prevent depolymerization back to monomer. Poly(dioxanone) has demonstrated no acute or toxic effects on implantation. The monofilament loses 50% of its initial breaking strength after 3 weeks and is absorbed within 6 months, providing an advantage over Dexon or other products for slow-healing wounds. Fibers made from polymers containing a high percentage of polyglycolide are too stiff for monofilament suture and thus are available only in braided form above the microsuture size range. The monomer p-dioxanone, is analogous to glycolide but yields a poly-(ether-ester) as shown below: Polydioxanone degradation in vitro was affected by gamma irradiation dosage but not substantially by the presence of enzymes. Poly(lactide-co-glycolide). Using the polyglycolide and poly (l-lactide) properties as a starting point, it is possible to copolymerize the two monomers to extend the range of homopolymer properties. Copolymers of glycolide with both l-lactide and dl-lactide have been developed for both device and drug delivery applications. It is important to note that there is not a linear relationship between the copolymer composition and the mechanical and degradation properties of the materials. For example, a copolymer of 50% glycolide and 50% dl-lactide degrades faster than either homopolymer. Copolymers of l-lactide with 25 70% glycolide are amorphous due to the disruption of the regularity of the polymer chain by the other monomer. A copolymer of 90% glycolide and 10% l-lactide was developed by Ethicon as an absorbable suture material under the trade name Vicryl. It absorbs within 3 4 months but has a slightly longer strength-retention time. Copolymers of glycolide with trimethylene carbonate (TMC), called polyglyconate, have been prepared as both sutures (Maxon, by Davis and Geck) and as tacks and screws Typically, these are prepared as A-B-A block copolymers in a 2:1 glycolide: TMC ratio, with a glycolide-tmc center block (B) and pure glycolide end blocks (A). These materials have better flexibility than pure PGA and are absorbed in approximately 7 months. Glycolide has also been polymerized with TMC and p-dioxanone to form a terpolymer 61