Biomaterials and their Medical Applications

Biomaterials and their Medical Applications

Article Type: Article

Authors: Mohammed Yaqob Shareef*

Abstract: Biocompatibility may therefore be  defined as the ability of a material to induce/perform with an appropriate host response in a specific application [1,2]. The response should be such as to ensure the continued safe and effective performance of the material [1,2]. Furthermore, such materials should not be toxic or have unfavourable properties either due to direct contact or breakdown products. In addition, factors including physical, mechanical properties, design, and surface chemistry should also be considered carefully.

Address: Ministry of Higher Education and Scientific Research, National University of Science and Technology, Thi Qar, Nasiriyah, Iraq. Corresponding author Email: myshareef62@yahoo.com

Keywords: Biomaterials, Biocompatibility; Physical and Mechanical Properties; Surface Chemistry; Materials/Tissue Interactions

Introduction

A biomaterial can be defined as any material, intended to interface with biological systems to treat, augment, or replace an organ or function of the body [1]. Any material used for this purpose, either in the oral cavity or another part of the body, therefore, has to be accepted by the bone and/or soft tissues. Biocompatibility may therefore be  defined as the ability of a material to induce/perform with an appropriate host response in a specific application [1,2]. The response should be such as to ensure the continued safe and effective performance of the material [1,2]. Furthermore, such materials should not be toxic or have unfavourable properties either due to direct contact or breakdown products. In addition, factors including physical, mechanical properties, design, and surface chemistry should also be considered carefully.

There are a number of terms which are commonly used in the description of materials/tissue interactions and are therefore important in the study of the biocompatibility [1]:

Bone bonding has been defined as the establishment, by physico-chemical processes, of continuity between implant and bone matrix [1].

Osteoconductive materials can only guide bone formation on their surface when implanted in a bony environment.

Osteoinductive materials have the ability to induce bone formation when implanted in non-osseous tissues.

Osseointegration a direct structural and functional connection between ordered, living bone and the surface of a load carrying implant, as assessed at the light microscopical level; modified from P-I Branemark, 1987 [1-3].

Interdigitation is the intimate association of the collagen fibres of the extra-cellular bone matrix with a materials surface, especially finger like projections into the material [4].

Bioactive a biomaterial that is designed to elicit or modulate biological activity [1,2]. It can be defined also as the ability of a material to actively influence the surrounding tissue and the development of a response from the tissue. In the case of bone, to aid the formation of a bone/material bond. Chemical bonding of bone with bioactive materials gives them their ‘direct’ attachment and interfacial strength which prevents interfacial fracture [5].

Bioinert materials that do not interact with the tissues of the body; forming no direct attachment with body tissues [5].

Biodegradation the breakdown of a material mediated by a biological system [1].

A wide variety of different materials have been used for medical and dental applications [6,7]. For example, cobalt-chromium-molybdenum (Co-Cr-Mo) based alloys are used in the production of subperiosteal implants as they are easy to cast [6,8]. They are highly resistant to corrosion and are the hardest and least ductile of all metals used in the field of dentistry. One of the advantages of metal implants is the necessary sterilisation process that can be achieved with conventional methods. Pure titanium and Titanium-6 Aluminium-4 Vanadium (Ti6Al4V) are currently the most commonly used metallic implants. Experimentally it has been shown that titanium has the potential to osseointegrate [9].

Carbon is a material exhibiting varied and unique properties not found in any other material. By controlling the structure of carbon, pyrolytic carbon implants were manufactured as a second generation material. Several different types of carbons are very inert and have been investigated as both dental and orthopeadic implants. But possibly the most useful application of these materials is in the vascular system. Carbons have a very good compatibility with blood, and many forms of these materials have been developed. Due to their unusual structure, they do not suffer from fatigue failure which can be a problem with other implant materials [10,11]. However, clinical complications reported following carbon implants are osteomyelities, parasthesia, anaesthesia, and substantial bone loss [10-12]. As a consequence their use today is limited.

In general, polymers are softer and more flexible and not as strong as the other materials described above including ceramics. Hence they are not particularly suitable for dental implant systems that require high mechanical stiffness and strength. They are generally used as graft materials [2,6].

There are two major type of bioceramics, namely bioactive and bioinert ceramics [10]. The bioactive ceramics include hydroxyapatite (HAp), tricalcium phosphate (TCP), glass ionomer cement (GIC) and bioglass [13-17]. When these types of bioceramic are used as bone substitutes, the bone is directly chemically bonded to the material. However, these materials are mechanically weak and unsuitable for use in stress bearing areas. Alumina and zirconia are classified as bioinert ceramics which are tough and strong [13,15,16]. When such ceramics are implanted as a bone substitute they become surrounded by a fibrous tissue capsule which prevents direct bonding of the material to bone. Repeated stimulation by stress and tension at the interface leads to the growth of connective tissue, which increases the risk of loosening between the bone and ceramic implant.

Hydroxyapatite (HAp) ceramic has an excellent affinity to the bone because hydroxyapatite and other calcium phosphate are also the mineral components of bone [13-16]. HAp has inadequate strength and toughness to be used in stress bearing areas [17]. Two approaches have been taken to overcome this problem. Ducheyne et al.[17] used HAp as a surface layer on titanium implants, but this approach has several problems that include separation of the coating layer from the underlying bioinert matrix and the coating process can reduce the strength of the bioinert materials [13]. The second approach is to produce ceramic composites with better mechanical properties [17,18]. The hypothesis is that the even distribution of the apatite phase as islets in the strong matrix phase or visa versa will contribute to mineralization and direct bone apposition onto this type of ceramic composite [18-20].

Bone

The production and implantation of bone substitutes requires understanding of the nature of bone. Bone can be defined as a  mineralised form of connective tissue composed of collagen fibres dispersed in a matrix of bone mineral with other constituents such as blood vessels and bone cells dispersed throughout. Bone can thus be regarded as a composite tissue, composed of inorganic and organic material. The inorganic component is mainly calcium and phosphate in the form of hydroxyapatite, whereas the organic component is mainly collagen and a little water. Collagen forms about 86% by weight of the organic matrix of bone [21-23].

For all its rigidity, bone is by no means a permanent and immutable tissue. Throughout its hard extracellular matrix are channels and cavities occupied by living cells, which account for about 15% of the weight of compact bone. These cells are engaged in an unceasing process of remodeling: one class of cells removes old bone matrix while another deposits new bone matrix. These mechanisms provides for continuous turnover and replacement of the matrix in the interior of the bone.

There are three main cells in bone [21-25] namely:

  1. osteoblasts, the bone-forming cells,
  2. osteoclasts, the bone-resorbing cells and
  3. osteocytes, which are of the connective-tissue cells family like fibroblasts and chondroblasts.

The bone matrix is secreted by osteoblasts that lie at the surface of the existing matrix into which it deposits fresh layers of bone. Some of the osteoblasts remain free at the surface, while others gradually become embedded in their own secretion. This freshly formed material (consisting chiefly of type I collagen) is called osteoid. It is rapidly converted into hard bone matrix by the deposition of calcium phosphate crystals. Once imprisoned in hard matrix, the original bone-forming cells, now called an osteocyte, has no opportunity to divide, although it continues to secrete further matrix in small quantities around itself. The osteocyte occupies a small cavity or lacuna in the matrix, but it is not isolated from other osteocytes. Tiny channels, or canaliculi, radiate from each lacuna and contain cell processes from the resident osteocyte, enabling it to form gap junctions with adjacent osteocytes [22,23]. Although the networks of osteocytes do not themselves secrete or erode substantial quantities of matrix, they probably play a part in controlling the activities of the cells.

While bone matrix is deposited by osteoblasts, it is eroded by osteoclasts. These large multinucleated cells originate, like macrophages, from hemopoietic stem cells in the bone marrow. The precursor cells are released as monocytes into the bloodstream and collect at sites of bone resorption, where they fuse to form the multinucleated osteoclasts, which cling to surfaces of the bone matrix and eat it away. Osteoclasts are capable of tunneling deep into the substance of compact bone, forming cavities that are then invaded by other cells. A blood capillary grows down the centre of such a tunnel, and the walls of the tunnel become lined with a layer of osteoblasts. To produce the plywood-like structure (known as haversian systems) of compact bone, these osteoblasts lay down concentric layers of  new bone, which gradually fill the cavity, leaving only a narrow canal surrounding the new blood vessel.

Bone growth takes place in one of two possible ways it may grow directly in fibrous connective tissue (intramembraneous ossification) or it may grow from cartilage firstly (endochondral ossification). Most bones show evidence of both these processes. Both processes continue until the bone achieves its natural size, then the endochondral ossification stops. Intramembraneous growth continues throughout life and contributes to the remodelling process which makes bone a dynamic structure. Bone can grow only by appositional (surface) growth (that is, by the laying down of additional matrix and cells on the free surfaces of the hard tissue), whereas young cartilage is capable of interstitial growth.

Mechanical Properties of Bone

Bone is a very dense, hard, specialized form of connective tissue. Like reinforced concrete, bone matrix is predominantly a mixture of tough fibres (collagen fibrils), which resist pulling forces, and solid particles (calcium phosphate as hydroxyapatite crystals) which resist compression. The volume occupied by the collagen is nearly equal to that occupied by the calcium phosphate, which is designed to provide mechanical support for skeletal motion and protect nonmineralized structures [22-24,26]. Its properties are a consequence of mineralized collagenous tissue structures that are continually remodelled in response to applied stess. Therefore, the mechanical properties of bone are dependent on the type of loading and the mineral content.

Collagen fibres in osteon units are unable to withstand compressive forces in the absence of hydroxyapatite that is deposited within the hole region and between fibres. The mineral composition of bony tissue varies from 30-90% for compact (cortical) bone and  5-50% for spongy (cancellous) bone. The elastic modulus is dependent on the porosity and increases with calcium content. The mechanical properties of bone were reviewed by Wasserman and Dunn [26] and some pertinent data are presented in the following table:

Table 1: Mechanical properties of long bone and skull

Site Modulus of Elasticity (GPa) Strength (MPa)
Tensile Compressive
Long bone 18 136 150
Skull 5.6 48 96

Cancellous bone, also called trabecular or spongy bone, is less dense than cortical bone and consequently has a lower modulus of elasticity and higher strain to failure, 5 to 7%. The difference in stiffness between the two types of natural bone tissues ensures a gradient in mechanical load across a bone.

Many types of implants for repair of the skeletal system are in contact with both trabecular and cortical bone. It is impossible for most prosthesis materials to produce similar gradients of stiffness between an implant and its host tissue.

The problem with a mismatch in elastic modulus across an interface is that the higher modulus implant (a limitation of bioinert ceramics is that materials with high fracture toughness and high flexural strength also tend to have a high elastic modulus and a low bioactive index)  will carry most of the load. The bone will be ”stress shielded”. which is undesirable because living bone must be under some tensile load to remain healthy. Bone that is unloaded or loaded in compression will undergo a biological change that leads to resorption and weakening. The interface between stress shielded bone and an implant will deteriorate as the bone structure is weakened.

All bioactive materials form a mechanically strong interfacial bond with bone, but most bioactive ceramics have a flexural strength and fracture toughness that is less than bone. Because of these biomechanical limitations, clinical use of bioactive ceramics has been restricted.

Healing of Bone

Following fracture of bone or implantation of material the trauma initially disrupts the normal blood circulation of the local bone and the adjoining soft tissue [21,22]. The local feeding vessels, which are the central capillaries passing through the haversian systems, are damaged. Blood from the damaged vessels flows out to surround the implanted material and fill any cavities in the bone and soft tissue. Osteocytes cut off from the nutrient supply die and the bone undergoes necrosis as far as the nearest unaffected canal. The degradation of tissue gives rise to free radicals which exert a weak bacteriocidal effect and stimulate fibroblasts to proliferate and attract macrophages. Osteoclasts invade the dead bone and resorb necrotic cells and bone matrix. Complimentary to this, osteogenic cells proliferate to begin the reconstruction of bone by elaboration of a non-mineralised collagenous matrix (osteoid), which subsequently undergoes calcification [22,23]. A material placed to act as a bone substitute should be designed to aid the above process and materials that exert a toxic effect on the cellular processes of healing/repair should be excluded from consideration.

Bioceramics

The potential of ceramics as biomaterials relies upon their compatibility with the physiological environment. Bioceramics are defined as ceramic biomaterials designed to achieve a specific physiological response when used as a material of construction for prosthetic devices or artificial internal organs. The compatibility of  bioceramics is the result of the fact that they can be composed of ions commonly found in the physiological environment, these are Calcium, Potassium, Magnesium, Sodium, and of ions showing limited toxicity to body tissue such as Aluminium and Titanium [27-29].

Ceramics fulfil many of the criteria of prosthetic materials. They are corrosion resistant, strong and have a high degree of biocompatibility. As a result, over the last decade ceramics have been increasingly studied with a view to their use as permanent implantable materials.

The first ceramics to be used as prosthetic materials were mainly alumina, titania and mixtures of calcia and alumina. These materials differ chemically from the composition of calcium phosphate ceramics, which may be why these materials did not integrate with the bone tissue [28,29].

Bioceramics are classified into three sub-groups, based upon their chemical reactivity in a physiological environment:

  1. Bioinert such as ZrO2 and Al2O3,
  2. Bioactive such as HAp and Bioglasses and
  3. Resorbable such as tricalcium phosphate, TCP, [27,28].

Bioinert Materials

Inert bioceramics or non-toxic inactive bioceramics undergo little or no chemical change during long-term exposure to the physiological environment. Even in those cases where the bioceramics may undergo some long-term chemical or mechanical degradation, the concentration of the degradation products in the adjacent tissue is easily controlled by the body’s natural regulatory mechanism [29].

Bone response to bioinert ceramics involves the formation of a very thin of fibrous membrane, several micrometers or less, surrounding the implant material. Bioinert ceramics are attached to the physiological system through mechanical interlocking, resulting from tissue ingrowth into undulating surfaces.

Alumina

One of the most widely employed bioinert ceramics is alumina (Al2O3). The first use of aluminas for implants in orthopaedics and dentistry was in the 1960s [30,31] and they were employed in hip prostheses as early as 1970 [31]. Since those early days the quality and performance of aluminas have improved [31-33] and high-purity, fine-grained aluminas are currently used for a wide range of medical applications, e.g. dental implants, middle ear implants, and hip and knee prostheses [31,32].

Although the aluminas currently available perform satisfactorily, a further improvement in strength and toughness would increase mechanical safety and may extent usage to higher stressed components. There are many reports comparing implant made of alumina ceramics to those made of stainless steel, on the basis of their resistance to corrosion or cytotoxity [27-29,34,35]. It is generally believed that alumina ceramics are superior to stainless steel in terms of the affinity of bone to these biomaterials in vivo [36-40]. However, alumina is sensitive to microstructural flaws which leads to a poor resistance to stress concentrations or mechanical impact in service [39,41].

Zirconia

Zirconia (ZrO2) is also classified as a bioinert ceramic [42] and is noteworthy for its high strength and toughness. Zirconia has three crystal phases: monoclinic, tetragonal, and cubic. Cubic-phase of zirconia is stable but brittle. Stress-induced transformation of the tetragonal phase to the monoclinic phase has been shown to increase the fracture toughness [42-45]. Through this phase transformation, zirconia has been shown to have a high fracture toughness. Phase transformation from tetragonal to monoclinic crystals, on the other hand, might result in a loss of strength with time [46].

In addition, zirconia has demonstrated better wear resistance than alumina and metal alloys in some experiments and is thus considered to be a desirable and promising joint material [45]. However, there is controversy about possible time-dependent deterioration in its mechanical properties, [47,48] which make orthopaedic surgeons hesitant about using zirconia [49]. Nevertheless zirconia, has become a material of great interest for medical and dental implant use as shown in biologic and histopathologic studies [50,51]. Partially stabilised zirconia has been shown to be tissue compatible, [51,52] and to possess twice the bending strength of polycrystalline alumina [52].

Bioactive Materials

In bioactive or non-toxic active bioceramics, the composition is designed such that the surface undergoes a selective chemical reaction with the physiological environment, resulting in a chemical bond between tissue and the implant surface [27-29]. The bonded interface protects the implant material from further deterioration with time. The main application of bioactive ceramics are in the area of ossicular bone replacement prostheses, as coating for orthopaedic applications and as dental implant materials.

Non-toxic resorbable bioceramics compositions contain only elements that are easily processed through normal metabolic pathways, such as calcium and phosphorous. The function of non-toxic, totally  resorbable bioceramics is to serve as a scaffold or filler of space permitting tissue infiltration and replacement. The main applications of non-toxic resorbable bioceramics is in the area of treatment of maxillofacial defects and for the oblitration of periodontal pockets.

Bioglass

The first man-made materials which were found to bond to living bone were glasses in system Na2O-CaO-SiO2-P2O5. These were discovered by Hench et al. [53] in the early 1970s and named Bioglasses, which can be classified as a bioactive material [54]. The bioglass material seems to enhance bone-implant interface development and bonding when implanted in bone [54-56]. A number of interfacial bonding mechanisms for bioactive glass substrata have been hypothesised by Hench et al.[53] and others [54-57]. These included mechanical interlocking and physical or chemical interactions, while similar mechanisms have been described for calcium ceramics by Jarcho et al.[58] when bioglasses are subjected to an aqueous environment, calcium and phosphate ions leach from the implant bulk to form a calcium phosphate-rich layer on the implant surface, that is believed to impart the bone bonding ability to the bioglasses. Biological processes, such as collagen interdigitation, have been described by Hench et al. [54,55,59-61] and possibly underlie bone bonding with surface reactive glasses.

Since the discovery of bioglass in early 1970s, various kinds of glasses and glass-ceramics have also been found to bond to living bone, and those are already clinically used such as an artificial middle ear bone [59] and alveolar ridge maintenance implants [62,63]. Bioglass is also used as artificial vertebrae, intervertebral spacers, iliac crest prostheses, and granules for bone defect fillers [61-63].

Glass Ionomer Cement

Glass ionomer cement (GIC) was orginally developed by Wilson and Kent in 1969  and is widely used as a dental restorative material. Biological evaluation has concentrated on the in vitro [64,65] and in vivo [66-68] responses of the dental pulp to GIC and compares favourably with other types of dental filling materials [64,69]. Brook [21] in his in vivo and in vitro study has confirmed that GICs showed a favourable biological response as potential bone substitutes, when compared to HAp and TCP ceramics. Brook reported also that GICs are a class of materials suitable for use as bone substitutes and cements and that GICs are bioactive materials releasing F-, Ca2+, PO42-, and SiO42-. Fluoride in a low systemic dose has been shown to stimulate bone formation [22,70]. The release of calcium ions may increase the degree of supersaturation of this ion in the tissue fluid adjacent to the implant and the silicate ion might provide favourable sites for the nucleation of apatite on the surface of the GIC which has been postulated to explain the bioactive response of  bioactive bioglasses to bone [71].

Calcium Phosphates

The fact that the mineral phase of bones and teeth is composed of calcium phosphate salts, has directed researchers to investigate these materials as potential candidates  for bone substitutes. Although the mineral is similar to hydroxyapatite (HAp), it is poorly crystalline in structure and contains a wide range of calcium phosphate phases, including TCP, carbonated apatite and numerous other ionic impurities such as fluoride, magnesium and sodium [72-74]. The development of synthetic calcium phosphates has provided a wide range of materials to be investigated for their bone replacement properties and may be the most biocompatible artificial bone substitutes currently used in the clinic. The common clinical applications of calcium phosphate ceramics are for jaw augmentation [74,75] and tooth root replacement [76], although their application is limited due to its poor biomechanical performance. These materials can bond with natural bone without fibrous tissue encapsulation [77-81]. The favourable response of the calcium phosphate ceramics, despite variations in the chemical composition and material structure, is attributable to their being mainly composed of calcium and phosphate ions [81,82].

Hydroxyapatite (HAp)

HAp with the chemical composition Ca10(PO4)6(OH)2 has been extensively study as a bone substitute. It shows good biocompatibility when implanted in either soft tissue [83-85], or hard tissue [82,86-88] and  can form strong and intimate bond with bone [85-87]. Driskell et al. [89] was the first to report that a chemical bond exists between bone and HAp. HAp, when sintered at high temperatures of up to 1250 oC, was initially believed to be non-resorbable [83]. However, it is currently generally accepted that degradation of HAp can occur to a certain extent [82,90]. LeGeros, et al. [91] reported that the bioactivity of  HAp may be related to its dissolution rate.

Fluorapatite (FA)

Fluorapatite Ca10(PO4)6(F)2 exhibits the same structure as HAp, the only difference being that the hydroxyl groups have been replaced by fluorine ions. The satisfactory biocompatibility of FA have been described by various investigators [92-94], while Lugscheider et al. [95] reported that FA is more stable than HAp at high temperatures, and more resistant to acid attack. It would therefore be more suitable than HAp for high temperature processes such as plasma spraying, which was confirmed by Dhert [94,96].

Tricalcium Phosphate (whitlockite)

Tricalcium phosphate Ca3(PO4)2 (TCP), together with HAp, is one of the most widely investigated calcium phosphate ceramics. TCP can be found in two forms, -TCP and the more stable -TCP. Both forms have been reported to exhibit relatively high dissolution properties [97-99]. TCP has been reported to possess good  biocompatible properties [97,98,100,101]. Klein et al. [100] reported  less bone contact and bone remodelling around plasma sprayed -TCP implants in dog femora, as opposed to plasma sprayed HAp and tetracalcium phosphate implants. Furthermore, it is thought to be less suitable as a plasma sprayed layer, due to its degradation character [94], and therefore is currently used primarly as a bone filler.

Tetracalcium Phosphate

Very little information is available for Ca4(PO4)2O (TTCP) as a bone substitute material. Similar interface morphologies as that with HAp were observed when TTCP was implanted in bone tissue [100,102]. Yoshimine, et al. [103] in their ultrastructural examination, showed that osteogenesis occurred directly on the surface of TTCP. This suggest that TTCP is biocompatible and possesses osteoconductive properties.

Future Direction and Applications

  1. Nanostrutured ceramics will be produced that demonstrate even higher degree of integration between biomaterials and the biological environment [104].
  2. To date, a large part of the interest has continued on titanium oxide (TiO2) nanotube because it is well known that titanium (Ti) is a biocompatible orthopaedic material which provides an excellent osseointegrative surface. Yet, little notice has been given to zirconium dioxide (ZrO2) nanotube [105].
  3. The mechanical properties have been found to be superior to those of the generally used Ti-6Al-4V alloy [106].
  4. There is a huge parametric space in which to explore novel nanostructures, nanopore dimensions and material compositions for improving and developing implant designs for anticipated tissue interactions.

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