Biomedical/bioengineering applications of carbon nanotube-based nanocomposites
This chapter discusses the use of carbon nanotube (CNT)-based nanocomposites for biomedical applications, particularly in the areas of joint replacement surgery and dentistry. The chapter initially reviews the current issues associated with orthopaedic implants and the application of CNTs to counterbalance some of the issues described. The chapter then discusses the use of CNTs in dentistry, including dental restorative materials, periodontal dentistry and denture base resins. The next section of this chapter discusses CNTs used in the areas of regenerative medicine and tissue engineering, including targeted drug delivery systems, monitoring biological systems and biosensors. The chapter concludes by discussing issues of CNT interaction with the body.
Materials used for medical devices or implants must meet a number of criteria ranging from physical, mechanical, biological, toxicological and other characteristics, depending on their particular application (Davis, 2003). One of the essential properties, biocompatibility, plays a vital role in the performance of the medical device when implanted in the body (in vivo) (Shard and Tomlins, 2006) and has been defined as the ‘ability of a material to perform with an appropriate host response in a specific application’ (Williams, 1999). The biological assessment of biomaterials and medical implants is governed by a set of standards developed by the International Standards Organisation (ISO) and known as ISO Standard 10993 or, in the United States (US), as US Food and Drug Administration (FDA) blue book memorandum #G95–12, which is an adaptation of ISO 10993. Moreover, the environment should not cause any degradation or breakdown of the biomaterial or medical device that would cause loss of its physical or mechanical properties, if not required, as with biodegradable biomaterials.
The development and manufacture of implantable medical devices have been ongoing since the early twentieth century. A common characteristic of materials used was their bio-inertness. The availability of biomaterials was a critical phase to meet the medical requirements of patients suffering from severe trauma or chronic debilitating diseases, for example, osteoarthritis, cardiovascular dysfunction and neuro-muscular disease. In the 1970s, research and development activities progressed from bioinert to bioconductive, bioactive and bioresorbable materials for biomedical applications. By the mid-1980s, it was possible to manufacture implantable devices, causing a biological response that could induce controlled reactions under in vivo conditions. By the end of the 1980s, the first bioactive materials had been developed for a range of musculoskeletal applications. Initially a discipline related to mechanical engineering, technological advancements in the medical device sector (for example functional biomaterials) have shifted boundaries and mind-sets. With tissue engineering entering the traditional sector of healthcare, the conventionally distinct borders between pharmaceuticals and medical devices are merging, and information technology has added new dimensions, presenting at the same time huge opportunities and difficulties.
Because of its significance, particular attention has been given to the orthopaedic sector, which focuses on rehabilitation and prevention of deformities, injuries and other disorders associated with the musculoskeletal system. Numerous artificial prostheses are available to stabilise bone fractures, spinal deformities or fractures, or to replace articular joints (Fig. 22.1) (Taylor, 2009). Several million people are affected by orthopaedic disease or trauma (Fig. 22.2) and the market size for orthopaedic implants is in the range of US $55 billion worldwide, with more than one million prostheses being implanted per annum (Phillips and Joshi, 2008).
22.1 Worldwide market for medical products (Taylor 2009).
22.2 Prevalence of osteoporotic fractures in the United Kingdom and United States (Phillips and Joshi 2008).
The annual growth rate for hip and knee implants is estimated to be 10% and 20–25% for spinal implants, due to the growing prevalence of musculoskeletal diseases (for example, osteoporosis and trauma-related fractures and osteoarthritis). Other major aspects causing growth are the ageing population with their related problems as well as injuries related to increased mobility and patient demands due to changing everyday lives.
The development of orthopaedic implants has evolved over many decades, resulting in improved devices with respect to reduced complications and improved longevity. Although the devices are successful, there still remain many unanswered problems and it is anticipated that with greater demand from patients for smaller and also bone conserving implants, new devices and biomaterials are required to prevent the current issues of implant failure, allergic reactions, lack of radiopacity, wear corrosion, degradation, etc. Approximately 10% of all surgical procedures involving implants are due to failure of the implant–tissue interface. The subsequent revision of surgical procedure is demanding and takes significantly longer; recovery takes many weeks and is less successful than the primary procedure. Moreover, the cost of revision surgery is significantly greater than the cost of the initial surgical operation. The reasons for failure of the implant are multi-factorial; however, the principal cause is aseptic loosening of the implants. Approximately, 75% of all implants fail due to mechanical breakdown of the implant–tissue interface in the absence of infection (Malchau et al., 2002).
Prosthetic implants and medical devices used for joint replacement surgical procedures are manufactured from almost all available material groups: metals for load-bearing applications like plates and nails for fracture fixation, rods for spinal fixation and prosthetic implants for total joint replacements. Structural ceramics find their use in articulation implants associated with total joint replacements, while bioactive ceramics based on calcium phosphate systems are used as bioconductive coatings and bone substitute materials. Polymers can be applied for articulation components of total hip joints and spinal implants, bone cements for stabilisation of prosthetic implants or for functioning as a temporary fixation device or scaffold, and in many other applications associated with repair of soft and hard tissue.
Although the success rate of hip joint (Fig. 22.3 (a)) and knee joint replacements (Fig. 22.3 (b)) is relatively high, a revision rate of 10% is anticipated within the first 10 years following implantation.
Loosening in the absence of infection (aseptic loosening) is the primary reason for failure and is usually a late complication often associated with an inflammatory response to wear particles from articulation of the joint replacements. Poor integration to the host bone as well as stress shielding are other reasons for failure at the implant–bone interface, resulting in aseptic loosening initially and the likelihood of peri-prosthetic fracture thereafter.
All the materials used as bearing surfaces in total joint replacement surgery are the same as, or minor modifications of, materials that have been used in clinical applications for many years. Ultra high molecular weight polyethylene (UHMWPE) has been used for the acetabular components in total hip arthroplasty, tibial components of total knee arthroplasty or the rotating component of artificial spinal discs since 1962. Metal on metal and ceramic on ceramic component combinations are in use for particular designs and primarily younger patients with more active lifestyles. All of the biomaterials currently used for joint replacement applications demonstrate limitations. Polyethylene (PE) is highly susceptible to wear, creep, degradation and fracture. Metals release ions into the surrounding tissue and bloodstream, which can cause systemic effects like hypersensitivity and potentially cancer. Ceramics are highly brittle and stress sensitive, therefore presenting specific loading risk and design limitations.
The anatomical stress pattern in bone is significantly changed after the implantation of a joint prosthesis or the stabilisation of a traumatic fracture. The implanted device will carry a portion of the load normally transferred through the bone, causing a change in the normal stress distribution. If the bone does not experience an appropriate load, then it will model Wolff’s law (Wolff, 1891), causing bone mass loss through resorption (atrophy) or the development of highly calcified bone in areas of high stress (hypertrophy). Metallic or ceramic-based implants exhibiting a Young’s modulus of 100–300 GPa are significantly stiffer than bone, which demonstrates a Young’s modulus of 8–24 GPa. Bone loss is associated with stress shielding from a modulus mismatch between the bone and the implant (Levenston et al., 1993) and can lead to implant loosening. This is not a serious issue when temporary devices are implanted for the stabilisation of a fracture caused by trauma; however, it is a problem for prosthetic implants that are used for replacements of diseased joins.
Nanotechnology presents openings not only to improve materials and medical devices, but also to develop novel smart devices and technologies like intelligent drug delivery systems (Webster et al., 2004). Several review papers discuss the different applications of nanomaterials and nanostructures for human healthcare (Salata, 2004; Liu and Webster, 2007). Artificial nanostructure and biomaterials have been postulated and studied for many medical applications to improve healing or replace tissue and organs (Lui et al., 2007).
Being hollow nanofibres, nanotubes offer interesting characteristics. Tubes are lighter than solid structures and offer potentially higher mechanical strength. As drug carriers, tubes have the capabilities to supply and release larger quantities than surfaces. Therefore, of all the nanofibres, nanotubes are of the greatest interest for many applications in medicine. Nanotubes, particularly short nanotubes, can be produced from several materials, potentially containing therapeutic drugs. Their surfaces can also be functionalised by attaching different functional groups, increasing their reactivity and hydrophilicity (Gojny et al., 2003). The derivative tubes demonstrate enhanced properties relating to ease of dispersion, solubility, management and processability. Functionalisation can also improve the interfacial bond to the matrix, enabling better stress transfer between the carbon nanotubes (CNTs) and the matrix (Eitan et al., 2006), and also making available potential sites for the attachment of chemical groups or therapeutic agents for more targeted delivery and efficacious application in medicine.
With their carbon composition, high aspect ratio, physical, mechanical and electrical properties, there has been significant attention in applying CNT technology to the medical and biomedical applications, for example, orthopaedic and dental implants and tissues engineered scaffolds. For several years, the number of research papers related to their use for biomedical applications has approximately doubled (Harrison and Atala, 2007). The biological activity and kinetics of any particular biomaterial depend on many variables, and the requirements and test protocols are well recognised for the conventional materials. However, because of the different physiochemical properties associated with their size, CNTs can potentially elicit a response in the human body that is very different and not directly expected from knowledge of the constituent chemical and compounds. For example, even a conventionally inert element such as gold can potentially be bioactive at the nanometre level (Goodman et al., 2004).
Too many factors and conditions at the nanometre level determine CNT interaction with cells and their superior structure for any planned use. As a consequence of this enhanced complexity, the biological assessment of CNTs is not fully understood at present, even for cell cultures or small animal models. In recent years many studies relating to the toxicity and biocompatibility of CNTs have been reported, thereby increasing our understanding of these materials under biological conditions (Fiorito et al., 2006).
CNTs are under investigation for potential application in many biomedical devices, given their inherent ability to interact with proteins and oligosaccharides, and their nanoscale dimensions that are comparable to those of basic biomolecules (Chen et al., 2001). CNTs have potential for applications in medicine, drug and delivery areas (Harutyunyan et al., 2002). There are many other possible applications of CNTs and their composites for medical applications, for example, MWCNT nanocomposites as biomimetic sensors, actuators and artificial muscles (Lee et al., 2005). The capacity to change the surface chemistry and properties through the addition of functional groups on the CNT backbone has led to potential applications ranging from vascular stents (Webster, 2007), platforms for neurone growth and regeneration (Wan et al., 2007), drug delivery vehicles for fighting cancer cells (Thuaire et al., 2004) and a possible protection of the immune system from bacteria and viruses (Bhargave, 1999).
When biomaterials come into contact with living cells, their surface morphology, structure, chemistry and charge dictate the biological host response. Because all biological processes, including traumatic or pathologic, are controlled and influenced by cell interactions at the nano-level (Chan et al., 2006), carbon nanostructures with their extraordinary properties have been proposed for several applications in direct contact with living tissue to allow specific cell attachment or rejection. One of the most exciting medical applications for CNTs and carbon nanofibres (CNFs), aside from drug delivery systems, is tissue engineering. This involves adapting and using biomaterials for building constructs that regenerate and repair functional tissue by influencing the organisation, differentiation and growth of the appropriate cells for body functions that have been damaged as a result of disease or trauma. CNTs and CNFs offer convincing properties for such applications, but their surface chemistry usually needs to be adapted to achieve the desired biological reaction (Harrison and Atala, 2007).
Tissue engineering is of particular interest in the healing of a traumatic or disease-related wound or defect, because the surrounding tissue has inadequate ability to bridge such distances. Without scaffolding, the defect would heal within the gap to form extended scar tissue that would compromise the structure, form and function of the repaired tissue. Scaffolds can be two-dimensional (e.g. soft tissue–skin) or three-dimensional (e.g. hard tissue–bone). 2D structures are usually achieved by surface medications, like incorporation of CNTs to substrates or as nano-filler in composites. The surfaces of different materials can be structured with nano-patterns by incorporating or applying CNTs for enhanced tissue–implant interaction. Equally important as the topography is the chemistry of the surface. 3D structures that allow transmission of all forms of load are normally built up by regular or irregular fibrous structures with interconnecting pores that allow the cells to penetrate and fill the scaffold. It has been demonstrated that MWCNTs can be shaped into 3D architectures and are ideal for cell seeding and in vitro cell modelling, leading to the design of exciting new tissue engineered scaffolds for biological applications. For example, the growth of a mouse fibroblast cell line on a 3D porous network based on an array of MWCNTs by exerting chemically induced capillary forces on the nanotubes was demonstrated by Correa-Duarte et al. (2004). The porous, interconnected network supported cell attachment and growth and demonstrated sufficient structural integrity to retain its shape in vivo, with adequate strength to support the developing tissue. Such a nano-structured material could serve as a biocompatible matrix to restore, maintain, or reinforce damaged/weakened tissues for applications where MWCNTs are required to act as drug delivery devices.
There is a clearly defined trend for medical devices to move from tissue substitution by artificial replacements to guiding and accelerating the healing process by means of tissue engineering, which can yield enormous personal and economic benefit. Because of their unique properties, CNTs can have a major role in the development of appropriate structures and surfaces. They have also been proposed recently for use in tissue engineered dental applications, delivered as an alginate nanocomposite gel at the surgical site (Kawaguchi et al., 2006).
Bone is a natural composite material, consisting of 10% water, 20% organic material and 70% mineral matter, by weight (Shi and Xuejun, 2006). The organic matter, which provides a framework consisting of mainly type-I collagen fibrils (Ø0.1–8 × 300 nm), is similar in scale to carbon nanotubes. The remainder of the organic material comprises other proteins, a cement-like substance, and a cellular component, consisting of osteocytes, osteoblasts and osteoclasts, which aid in dissolution, deposition and nourishment of the bone. The inorganic mineral component, which provides strength is a calcium-deficient, carbonate-substituted apatite, containing calcium and phosphate ions, similar in structure and composition to hydroxyapatite (Ca10(PO4)6(OH)2) (LeGeros and LeGeros, 1993).
Figure 22.4 illustrates the hierarchical structure of bone. Bone mineral crystals (Ø 2–5 × 20–50 nm) are arranged between the ends of collagen fibrils, which are then rearranged into sheets called lamellae. The morphology of mineral crystals is generally agreed to be plate-like with dimensions on the scale of ten to hundreds of angstroms. The collagen lamellae are arranged either in concentric circles called tubular Haversian systems or in sheets. These sheets form the sponge-like cancellous or trabecular bone found inside the structure at the bone ends. The Haversian system configuration, on the other hand, leads to dense, cortical form of bone, which comprises 80% of bone mass and surrounds the cancellous bone (Lakes, 1993).
Because tissue development is controlled by events at cellular and molecular level, the surfaces for osteointegration need to be able to influence osteoprogenitor population activity and function. Recent review papers have addressed the similarities between bone and carbon nano-structures and report the application of CNTs and CNFs (Christenson et al., 2006; Webster and Ah, 2006; Liu et al., 2007). Significant research is currently focussed on looking at strategies for developing biomimetic tissue surfaces for prostheses; surfaces that have been engineered at the nano-scale to mimic or interact with soft or hard tissue, thereby behaving as a living surface. With the potential that nano-technology offers, many current problems in developing long-lasting, biocompatible orthopaedic devices might be solved in the future. Several applications for nanostructures and nanomaterials are already being pursued and early results are encouraging (Streicher et al., 2006). The application of CNFs and CNTs has been proposed and investigated as a standalone product, either functionalised as a scaffold or as reinforcement for polymers or other composite materials for different orthopaedic applications, mostly focusing on hard tissue repair (Balasundaram and Webster, 2006; Christenson et al., 2006; Webster and Ah, 2006).
Price et al. (2004) compared polymer casts of consolidated typically sized and nano-sized carbon fibres, highlighting the importance of the nanostructure. This study showed significantly enhanced osteoblast adhesion to the nano-phase fibres. In a further study by Price et al. (2003), they produced composites with similar fibres using a polycarbonate–urethane polymer matrix. Such composites have already attracted significant attention for orthopaedic applications because of their tailored electrical and mechanical properties through the incorporation of CNFs. In these studies, Price et al. (2003) showed the advantage of adding 10 to 25 wt% CNFs for osteoblasts. Research groups have demonstrated selective adhesion of osteoblasts on composite structures containing CNFs, therefore making them potential candidates for prosthetic implants (Elias et al., 2002). They incorporated CNFs and CNTs in a polycarbonate–urethane polymer matrix to create a nanostructure added to the structure, thereby mimicking those nanofeatures evident in living tissue, and assessed these nanocomposites for efficacy using cell lines in vitro. Selective adhesion was more pronounced with increased content of functionalised CNT–CNF; osteoblasts exhibited increased viability (Elias et al., 2002; Webster et al., 2004) and were preferred over fibroblasts (Price et al., 2004). In another study by Khang et al. (2005) using the same nanocomposites, it was demonstrated that fibronection adsorption also increased with increasing quantities of MWCNTs, a consequence of the increased surface roughness and surface energy of the nanocomposite. Osteoblasts align along CNFs with a pyrolytic outer layer, that is reduced surface energy and depose calcium phosphate mineral, which is an indicator of their viability (Khang et al., 2006). All of these phenomena would reduce fibrous tissue formation around a prosthetic implant and are advantageous for optimal osseo-integration.
Shi et al. (2005) developed an injectable nanocomposite based on polypropylene fumarate containing unfunctionalised and functionalised SWCNTs as a potential candidate for load-bearing orthopaedic applications. They concluded from their investigations that the functionalised SWCNTs were more homogeneously dispersed within the matrix, which improved the mechanical function of the nanocomposite even at levels of loading as low as 0.1 wt% SWCNTs.
Functionalised SWCNTs have also been applied to developing hydroxyapatite (HA) based bone substitute materials and tissue promotion scaffolds for the treatment of bone fractures (Zhao et al., 2005). The methodology amalgamates the osteo-conductive properties of HA with the supra-mechanical properties of CNTs, thereby improving stiffness and strength of the brittle ceramic-based bone substitute material. For this purpose, SWCNTs were chemically functionalised with various chemical groups to produce negative surface charges that attract calcium ions from the surrounding electrolyte. Some of these conjugates led to nucleation, self-assembly and orientation of the HA crystals, facilitating control of their alignment; while other groups enhanced biocompatibility of the CNTs, improving their solubility in water. This research has potential application to artificial bone with improved flexibility and strength, novel types of bone grafts and a possible local treatment of osteoporosis. This could be achieved by delivering an aqueous solution of functionalised CNTs into the bone fracture or defect, encouraging new bone tissue to grow and fill the defect. Zanello et al. demonstrated under laboratory conditions that osteoblast cells can grow and proliferate best on scaffolds comprising SWCNTs and MWCNTs, when the functionalisation of these CNTs can result in a neutral surface charge (Fig. 22.5) (Zanello et al., 2006).
Another technique for improving the durability and biological acceptance of biomaterial surfaces, for applications as orthopaedic prostheses, is the application of coatings consisting of HA and CNTs. Balani et al. (2007) reported the improvement in the fracture toughness of a plasma-sprayed HA coating by incorporating 4 wt% MWCNTs into the HA. Additionally, they observed no adverse cellular response when in vitro biocompatibility studies were conducted on the HA–MWCNT coating. White et al. (2007) presented an overview of biomedical applications in the context of HA–CNT nanocomposites.
Another CNT-based nanocomposite for biomedical application can be manufactured by blending MWCNTs and poly-L-lactide (PLLA). The interaction between the polymer matrix and the MWCNTs takes place mainly through the hydrophobic C-CH3 functional groups (Zhang et al., 2006a). The conductivity of the MWCNT–PLLA nanocomposite improved as the level of MWCNT loading increased and the growth of fibroblast cells was inhibited. Potential applications of such a nanocomposite could be as a biodegradable, conductive material with selective cell interaction facility as required. Nonetheless, the important question about the fate and the biological reaction of the released CNTs from a resorbable polymer matrix needs to be fully understood.
Electrospinning can be used to produce fibres containing CNTs, for example, SWCNT-reinforced silk nanofibres (Ayutsede et al., 2006) and incorporating as little as 0.8 wt% of MWCNTs to a chitosan derived biopolymer increased its tensile strength by 200% (Wang et al., 2005). MacDonald et al. incorporated 4 wt% SWCNTs into collagen, achieving an electrically conductive composite with improved mechanical properties (MacDonald et al., 2005; MacDonald and Stegemann, 2006). Such a CNT-collagen-based nanocomposite could potentially have an application as a scaffold material in tissue engineering. Correa-Duarte et al. applied chemically induced capillary forces to interconnect MWCNTs to a 3D network, which allowed for the successful growing of murine derived fibroblast cells within the 3D structure (Correa-Duarte et al., 2004). Firowska et al. (2006) merged lithographic and layer-by-layer self-assembly technologies to produce a highly orientated 3D scaffold structure based on inter-crossed MWCNTs for potential application in tissue engineering.
A strong acetabular socket prosthesis, which is thin and compliant, but could replicate the normal transfer of load-bearing forces to the supporting bone should significantly decrease bone atrophy (loss) and provide long-term joint stability. Initial efforts using a composite based prosthesis consisting of a flexible horseshoe-shaped artificial acetabular socket were developed (Jones et al., 2001a). Jones et al. (2001b) reported the efficacy of a polybutylene terephthalate reinforced with carbon fibre (CF) prosthesis in 50 patients over the age of 80 years after five years implantation. The initial results were positive in terms of performance; however, improvements could be made to the design. For example, if a mono-block composite acetabular socket could be developed that encourages new bone growth on the outer surface of the prosthesis and is, in parallel wear-resistant, then this would be an exciting prospect. Analogous possibilities are available, taking into account current developments for segmental and partial joint replacement. For example, resurfacing of the hip joint, which is presently achieved using large diameter metal-on-metal articulations for young and active patients. There is currently significant interest to maintain as much natural bone as possible to facilitate bone preservation for any revision surgical procedures, as the life expectancy of the younger patient population is greater than the working life of the prosthetic implants. The pursuit of smaller implants also challenges the limits for the existing materials, and there is a need for new materials and technologies to meet the requirements of such designs. Several applications for trauma treatment, spinal repair and the treatment of degenerative diseases also require development. The application of CNTs could achieve the objective of such developments, especially when adopting all properties and not just the mechanical characteristics, and could have a major impact on new designs and implant prosthesis.
To improve the mechanical strength and resistance of prosthetic implants, the reinforcement of high strength polymers with CNTs appears to be a real prospect. CNTs are highly anisotropic and due to their high strength, high aspect ratio and excellent thermal and electrical conductivity, they are currently been used as an alternative to CFs for producing composite materials. CNTs have generated significant interest, particularly as a reinforcing material for polymer matrix composite materials. CNTs are believed to improve the mechanical, electrical and thermal properties of those composites. Nevertheless, several complications related to CNT dispersion, their compatibility within the polymer matrix and the nature of the bond between the CNT and matrix material need to be understood before practical applications. To develop the unique properties of CNTs in a polymer composite, fundamental challenges need to be tackled that have proved to be a difficult issue with a number of opposing variables: including CNT production and purification, functionalisation of CNTs, alignment and dispersion of CNTs, and processing method used (Mylvaganam and Zhang, 2007; Weisenberger et al., 2007).
A strong interfacial bond between the CNT and polymer matrix is required to facilitate the appropriate stress transfer between the reinforcement and matrix phases of the nanocomposite (Lau Kin-Tak, 2002). Three main load transfer mechanisms control the principal stress transfer: (1) surface topography and micromechanical interlock; (2) chemical bonding; and (3) van der Waals interaction. To achieve bonding between the reinforcement and the polymer matrix two approaches have been adopted: (1) non-covalent attachment of molecules, resulting in the formation of weak van der Waals forces; and (2) covalent attachment of functional groups to the CNT walls. The non-covalent approach has the benefit that the perfect structure of the CNT is not altered, thus their overall mechanical properties remain unchanged. The covalent attachment of the functional groups to the CNT surface can be achieved by many chemical or physical methodologies. Adopting one of these processes has the potential to introduce defects on the CNT walls and therefore reduce the mechanical strength of the reinforcing material. It has also been demonstrated to have an influence on the stiffness of the CNT (Namilae et al., 2004), which needs to be considered. Notwithstanding this fact, it is thought that the interfacial shear strength of the polymer–CNT nanocomposite with non-covalent bonding can be increased over an order of magnitude by incorporating < 1 wt% chemical bonds between the CNT and the polymer matrix. Another aspect is the effect of any reinforcing material on the structure of the polymer matrix. The addition of any additive during the processing of a thermoplastic semi-crystalline polymer can potentially influence the crystallinity and therefore the final properties of the polymer–CNT nanocomposite. The trade-off between strength of the polymer–CNT interface, CNT strength and the final properties of the nanocomposite material must be well balanced (Qian et al., 2000).
The majority of studies have focused on the development of polymer–CNT nanocomposites that could be applied as lightweight, high strength, fibre-reinforced materials for non-medical applications, showing that, apart from the polymer, the CNTs and any potentially applied functionalisation, there is also a balance between the level of CNT loading required to achieve a significant improvement in the properties of the nanocomposite and its ability to be processed (Khare and Bose, 2005). Many different techniques have been used to attain a strong interfacial bond between various polymer matrices and the CNTs. SWCNTs and MWCNTs have been used for reinforcing thermosetting and thermoplastic based polymers (Thostenson et al., 2005). Ami Eitan et al. (2003) have shown functionalisation of SWCNTs and MWCNTs by using a solution of nitric and sulphuric acid to form carboxylic acid groups on the surface, which are formed along the CNT walls and at the end-caps. This approach seems to be ideal for optimal transfer within the composite material. They reported that it is possible to further react with the epoxide functional group to facilitate even better interaction between the polymer matrix chains and the CNT surface of the composite.
Only a few studies to date have explored the potential use of CNT composites for load-bearing orthopaedic applications. In an effort to enhance the mechanical performance of UHMWPE, a well-accepted hydrophobic and non-polar articulation material for hip and knee joint replacements, by incorporating CNTs to the polymer matrix, reinforced composites have been manufactured by research groups from Sweden (Emami, 2007) and France (Babaa et al., 2006). The French group used UHMWPE with unfunctionalised SWCNTs as a reference, which resulted in the non-covalent attachment of CNTs and the polymer matrix.
Although in this approach the structure of the SWCNTs is not changed, the mechanical properties of the nanocomposite would only be affected minimally and the low attraction forces between polymer matrix and CNTs are unfavourable (Holzinger et al., 2004), as the clinical history of the same material has demonstrated (Pryor et al., 1992). Babaa et al. (2006) also covalently bonded CNTs and polymer matrix with chemically modified SWCNTs and polyethylene (PE). This was achieved by oxidising the CNTs, forming carboxylic acid groups along the CNT walls and end-caps, and thermally oxidising the UHMWPE, forming oxygen-containing functional groups and using a di-amine to bond the oxidised PE and SWCNTs. The bi-functional molecule can react with the carboxylic acids groups attached on both the SWCNTs and PE to form a strong covalent bond. The resulting nanocomposite was formed as a film and fibres. Mechanical testing of the nanocomposite material showed significant improvements over the reference material. Another interesting aspect of applying CNTs to reinforce UHMWPE is that MWCNTs can act as radical scavengers and antioxidants (Fearon et al., 2002), especially in view of the fact that the medical grade of UHMWPE is only available without any antioxidant, which tends to degrade when implanted in the body.
PEEK, a semi-crystalline high strength polymer, has been studied for many nanocomposite applications. Sandler et al. (2002) reported on the mechanical properties of a vapour grown 150 nm Ø CNT reinforced PEEK composite using extrusion technology to produce masterbatches. The final specimens were produced by using the injection moulding process. A linear increase in stiffness and yield strength with loading concentrations of up to 15 wt% was demonstrated, while PEEK matrix ductility was sustained to 10 wt% CNT loading. No influence of reinforcing material on the polymer matrix crystallinity was measured. Werner et al. (2004) studied the wear properties of PEEK–CNT nanocomposite. An exponential wear rate reduction for composites with 5–10 wt% CNT was demonstrated, depending slightly on the reinforcing material content. Werner et al. postulated that the CNTs demonstrated the properties of a solid lubricant, therefore decreasing the wear rate.
With the intention of manufacturing load-bearing orthopaedic prosthetic implants, Babaa et al. (2006) also used PEEK as a polymer matrix material and reported production of UHMWPE–CNT and PEEK–CNT nanocomposites. To obtain a covalent bond between the components, the CNTs were functionalised by using di-functional molecules for a better bonding between CNTs and the polymer matrix. With this methodology, the properties of the nanocomposite were enhanced. Bantignie (2007) attached functionalised MWCNTs and sulphonated using direct reaction bonding to sulphur group of PEEK, and demonstrated using spectroscopic methods that a covalent bonding between SWCNTs and the PEEK polymer matrix could be achieved. The direct attachment reaction bonding method is an unwieldy, labour intensive process, with limited commercial applicability. However, the functionalised CNTs have been shown to achieve a high level of bonding with functionalised PEEK, a clinically applicable material.
The premise of a PEEK–CNT nanocomposite material offers an excellent opportunity for outstanding applications of thin and high strength prosthetic implants with exceptional wear properties that could also be engineered on their interface with the host tissue to be fully integrated.
Another biomaterial used in joint replacement surgery is PMMA, which is the primary component of acrylic bone cement. It is well documented that PMMA cement is susceptible to fatigue-related cracking and impact-induced failure (Kuehn et al., 2005). Active or overweight patients with implants fixed with PMMA cement are at risk from cement mantle failure, which occurs in 5% of all total joint replacements patients post-operatively. Failure rates of 67% have been reported after 16 years in patients younger than 45 years. It is postulated that the annual number of revision total knee arthroplasties performed in the US will increase 601% to 270,000 by 2030 and the number of hip arthroplasty revisions will increase 137% to 97,000, with the main cause of implant failure attributed to cement mantle failure (Marrs et al., 2006). There is a plethora of literature reporting methods for improving the mechanical and physical properties of PMMA bone cement. Many of these studies have incorporated various additives into the polymer matrix with the aim of improving the mechanical properties. Specifically, PE fibres (Uzun et al., 1999), glass fibres (Stipho, 1998), long macroscopic carbon fibres (Manley et al., 1979), and titanium fibres (Topoleski et al., 2004) have all been employed in attempts to bridge fatigue cracks and prevent, or reduce, the rate of their propagation. The involvement of these additives have been less than successful due to the poor fibre–polymer matrix bonding (and subsequent debonding), increased viscosity, poor additive distribution, and the adverse effects such materials have on the mixing of bone cement (Marrs, 2007). Marrs et al. (2006) investigated the influence of MWCNTs in PMMA-based bone cements. They reported moderate improvements (13–24%) in the static properties when 2 wt% MWCNTs were incorporated into the methyl methacrylate-styrene cement. Marrs et al. (2007) also reported significant improvements (> 300%) in the dynamic properties when MWCNTs (2 wt%) were added to the same bone cement. However, both studies (Marrs et al., 2006; Marrs et al., 2007) used non-clinically relevant methods to ensure optimal dispersion of the MWCNTs into the cement. Additionally, it has been hypothesised that functionalisation of the MWCNTs could potentially improve the dispersion and interfacial bonding of the MWCNTs within the cement matrix, leading to further improvements in the static and dynamic properties (Gojny et al., 2003; Marrs et al., 2007). It has been reported that the static properties of PMMA polymer resin can be significantly improved through functionalisation of MWCNTs (Pande et al., 2009). A subsequent study by Ormsby et al. (2010a) incorporated 0.1 wt% MWCNTS (unfunctionalised and carboxyl functionalised) into PMMA bone cement and determined the mechanical and thermal properties. They reported the extent of the effect was dictated by the type of MWCNT, method of introduction used and the properties being quantified. Improvements in mechanical properties (2–32%) were attributed to the MWCNTs being well dispersed within the PMMA cement when the MWCNT were ultrasonically disintegrated in the liquid monomer prior to mixing with the PMMA powder, thereby arresting/retarding crack propagation through the cement (Fig. 22.6 (a)). Conversely, reductions in mechanical properties were ascribed to MWCNT agglomerations occurring within the cement microstructure when the MWCNTs were dry blended into the PMMA powder prior to mixing with the liquid monomer (Fig. 22.6b). The degree of these agglomerations was dependent on the method used to incorporate the MWCNTs into the cement. Ormsby et al. (2010a) also reported that MWCNT-reinforced PMMA cement shrinks less and demonstrates a lower exothermic thermal reaction. The level of heat produced within the exothermic polymerisation reaction of the PMMA bone cement was significantly reduced when functionalised MWCNTs were added. The maximum temperature value reduced by 29–53% when compared to the control cement, as measured in accordance with ISO 5833: 2002. However, this reduction in maximum temperature extended the duration of the polymerisation reaction by 52–82%. Notwithstanding this fact, the reduction in exotherm translated in a decrease in the thermal necrosis index value, which is indicative of a potential reduction in hyperthermia experienced in vivo by polymerisation reaction of PMMA cement.
22.6 (a) 0.1 wt% carboxyl functionalised MWCNT–PMMA cement in which MWCNT can be seen to bridge a micro-crack on the surface; (b) 1.0 wt% unfunctionalised MWCNT–PMMA cement showing an agglomeration of MWCNTs (indicated by arrows), which was the fracture initiation point for this specimen.
A further study by Ormsby et al. (2010b) investigated additional mechanical augmentation of PMMA bone cement via the incorporation of MWCNTs with varying functional groups (carboxyl and amine) at increasing wt% loadings (up to 1 wt%). The study concluded that incorporating low loadings of MWCNTs (≤ 0.25 wt%) into PMMA bone cement improved the mechanical properties of the resultant nanocomposite. Significant improvements in compressive strength (≈ 40%), bending strength (≈ 18%), compressive modulus (≈ 28%), bending modulus (≈ 14%) and fracture toughness (≈ 61%) were observed when carboxyl functionalised MWCNTs were added to the PMMA bone cement. Higher loadings (≥ 0.5wt%) provided lesser improvements in the mechanical properties, and in some cases significant reductions were recorded. The highest loading of carboxyl functionalised MWCNTs (1 wt%) significantly reduced the mechanical properties of the PMMA bone cement, giving reductions in compressive strength (≈ 26%), bending modulus (≈ 42%), compressive modulus (≈ 9%) and bending strength (≈ 44%), respectively. The incorporation of amine functionalised MWCNTs followed a similar trend with the 0.1 wt% loading giving the maximum improvement in mechanical properties.
The dynamic mechanical behaviour of MWCNT–PMMA composites has also been studied (Jin et al., 1998; Jin et al., 2001; Wang et al., 2006). Wang et al. (2006) showed that the storage modulus of PMMA at room temperature is doubled upon addition of 26 wt% MWCNT. Subsequent studies have used smaller amounts of functionalised MWCNTs leading to significant improvements in the storage modulus of PMMA. Velasco-Santos et al. found that the storage modulus of PMMA at 40 °C increased by 50% upon addition of 1 wt% functionalised MWCNTs and by 66% upon incorporation of 2 wt% functionalised MWCNTs (Velasco-Santos et al., 2003). Hwang et al. (2004) reported that the storage modulus of PMMA increased by 1100% when 20 wt% of PMMA-grafted MWCNTs were added. Building on previous studies using general purpose PMMA as the matrix material, Nien and Huang (2010) incorporated up to 0.75 wt% MWCNT into PMMA bone cement and determined the effects on the static and dynamic mechanical properties. They reported an improvement in the compressive strength (21%) and tensile strength (15%) of the PMMA cement when 0.75 wt% MWCNTs were added. Using dynamic mechanical analysis, they observed that the glass transition of the PMMA cement reduced by 8% for the same PMMA–MWCNT cement combination (Nien and Huang, 2010). They postulated the reduction in glass transition temperature was due to the CNT acting as a plasticiser. Improvements in the storage modulus were not demonstrated for the PMMA–MWCNT cement, on average a 7% reduction was observed irrespective of the wt% MWCNT added to the PMMA cement.
The incorporation of CNTs in PMMA cement has a significant effect in avoiding crack propagation, and could deal with one of the primary reasons for revision surgery of a cemented total joint replacement: fracture and fatigue fracture of the PMMA cement mantle. Advantageous side effects are that such reinforced PMMA cement shrinks less and demonstrates a lower exothermic thermal reaction, and that the hollow CNTs would facilitate the addition of pharmaceutical components like antibiotics, anti-inflammatory drugs and chemotherapeutical agents, etc. Potential causes for consideration may be the increase in the rheological properties and setting characteristic of such a PMMA–CNT bone cement system during application, and a possible phase separation between the PMMA cement and CNTs while filling the space between the prosthesis implant and the bone. Another potential consequence that needs to be investigated is the interaction between the CNTs and the metal-based prosthetic implant with respect to fretting, abrasion and corrosion, particularly at higher levels of CNT loading.
Alumina–CNT nanocomposites have many potential medical applications, particularly in orthopaedic prosthetic implants. Improving the low toughness, poor bending strength of alumina by enhancing the resistance to crack propagation would allow for greater design flexibility, less risk of fracture and better wear resistance. Such improvements would be advantageous for implants manufactured from alumina, which currently have several disadvantages due to fracture risk.
Many studies have reported improvements in the properties of metal and ceramics by producing metal–CNT and ceramic–CNT nanocomposites (Zhan et al., 2003; Thostenson et al., 2005; Zhang et al., 2006b). As a brittle material, alumina is strong in compression but is prone to fracture, that is poor flexural strength and fracture toughness. The alumina quality used for artificial hip and knee joint bearings has a typical flexural strength of 500 MPa and a fracture toughness of 4 MPa m1/2. To improve the resistance to crack initiation and propagation in the microstructure of the alumina, either fibres or toughening constituents that can absorb components of the strain energy induced by the crack are employed. CNTs could have such an influence at the nanoscale level. To produce alumina–CNT nanocomposites, various approaches have been put forward (Zhan et al., 2003), however, limited data is reported on the effects of incorporating CNTs into the alumina. Notwithstanding this fact, experience thus far has demonstrated difficulty in obtaining a dense and a well-dispersed structure alumina–CNT nanocomposite material.
In spite of this, Zhan et al. (2003) reported the successful manufacture of dense nanocrystalline alumina–SWCNT nanocomposites at sintering temperatures as low as 1150 °C by spark plasma sintering, and achieving a fracture toughness of 4 MPa m1/2, nearly three times that of pure nanocrystalline alumina. Conventional alumina ceramics are sintered at temperatures above 1400 °C in an oxidising atmosphere that would burn the embedded CNTs. Therefore, the nanocomposites have to be sintered at much lower temperatures (1000 °C). It is highly unlikely that oxide ceramics meeting today’s strict standards can be manufactured in this way. A possible solution to a lower sintering temperature is to allow more glassy phases between the ceramic grains, for example, by using a less pure alumina; however, this reduces the overall quality. A more promising approach to higher strength ceramics might be the application of nanostructure constituents (Tanaka et al., 2002).
Dental materials have to be high performance materials, to survive within the complex and hostile oral environment. They will experience a wide range of temperatures (15–68 °C), pH (6.0–7.4), chemical exposure from food and drink, together with static and fatigue loads of up to 600 N. It is therefore essential that all restorative dental materials, whether they are to be used to restore a cavity in a tooth, or make a fixed dental prosthesis, such as a crown or bridge, or a removable dental prosthesis, such as a denture, are able to give optimum performance over a period of ideally 8–10 years before replacement is required.
CNTs have received a modest amount of research attention in dentistry as they have shown potential to increase the strength of composite materials and implants, increase cell adhesion and proliferation, effect nucleation of hydroxyapatite and provide protection against bacteria (Akasaka et al., 2009).
However, the main obstacle to the widespread inclusion of CNTs in dental materials is the deep black colour that they bestow on any composite material to which they are added. Obviously, black is not going to be the colour of choice for a dental restoration, as patients require aesthetic materials that blend in with the natural colour of their teeth and gums. However, if the improvement in mechanical and physical properties is sufficiently high, it may be cost-effective to develop techniques and procedures to mask the black colour of the CNTs. This could be carried out using an aesthetically coloured coating of approximately 1–1.5 mm thickness, or by the incorporation within the composite material of other inorganic additives such as sol-gel-based opalescent fillers, or chromophoric xerogel pigment particles (Zhang et al., 2008).
It is important when restoring a cavity within a tooth that the replacement material bonds to the surface of the tooth to prevent microleakage of bacteria, found in saliva, between the filling material and the cavity wall. If this occurs, as it may do with current dental materials, the bacteria can enter the dental pulp, containing the nervous and vascular tissues, and cause inflammation. The inflammation in turn causes pain, and may lead to death of the pulp tissues, necessitating either tooth extraction or root canal treatment.
Sophisticated mechanisms of micro-mechanical and chemical adhesion have been developed to try and prevent this microleakage. These involve the application of acids to the surface of the tooth dentine, which demineralise the dentine, leaving collagen fibres exposed, and stripping back the smear layer exposing the dentinal tubules. Dental adhesive materials are then applied which flow into the openings of the tubules and inter-mingle with the collagen fibres creating a micro-mechanical bond. This method works satisfactorily, however, it is still vulnerable to microleakage in the long term. Therefore, constant research is being carried out to determine if the bond to tooth substances can be improved.
In early studies, CNTs were applied to tooth dentine and cementum and showed that they selectively adhered to these dental surfaces, possibly by adhering to the exposed collagen fibres (Akasaka et al., 2009). Dentine contains small, thin apatite crystals embedded in a protein matrix of cross-linked collagen fibrils (Nicholson, 2006). The interaction between the CNTs and collagen was investigated using SEM and fluorescence confocal laser scanning microscopy. Microscopy showed that the CNTs adhered to the surface of tooth dentine and cementum, but could not penetrate into the dentinal tubules. There was no interaction observed with tooth enamel, which does not contain collagen. In addition, the tooth surfaces to which the CNTs attached were observed to change colour to grey, indicating the presence of CNTs.
These findings are in agreement with those of earlier papers which also reported strong interactions between CNTs and collagen fibres in an aqueous environment (MacDonald et al., 2005; Liao et al., 2007). They speculated that as proteins in solution are known to adsorb CNTs via hydrophobic interactions, which the CNTs were probably interacting with the collagen fibres using the same mechanism (Karajanagi et al., 2006).
Further studies by Akasaka et al. (2009) showed that the bond strength between the CNT-coated dentine and a resin composite dental filling material was not decreased by the presence of the CNTs. The authors speculated that further work in the area of dental bonding was warranted as CNTs can exert beneficial effects such as the nucleation of hydroxyapatite and protection against dental bacteria without decreasing the bond strength to restorative materials. The ability to protect against dental bacteria would be of particular benefit in counteracting the effects of any microleakage.
CNTs have been added to dental methacrylate-based resin composites used to restore cavities in teeth (Zhang et al., 2008). These dental filling materials are conventionally tooth-coloured and therefore highly aesthetic and in demand by patients. However, although many significant advances have been made in the wear and physical properties, together with the dimensional stability of these materials over the last 15 years, the average life expectancy of a resin composite restoration is significantly lower than that of a conventional dental amalgam restoration (Letzel, 1989; Mjor et al., 1990; Chadwick et al., 2002; Lucarotti et al., 2005).
Efforts therefore continue to improve, in particular, the physical properties of dental resin composites, often by the addition of nanoparticles. The use of nanoparticles within dental resin composites is increasing in popularity as the small size and wide size distribution can achieve increased filler loading, giving improved wear and reduced shrinkage on photo-polymerisation. Increases in tensile and compressive strength and fracture toughness have also been reported (Baseren, 2004; Yap et al., 2004a; Yap et al., 2004b; Chen et al., 2006; Mesquita et al., 2006).
The majority of research has focussed on the addition of tooth-coloured or white nanoparticles, such as nano-hydroxyapatite, colloidal silica, titanium dioxide or aluminium oxide nanoparticles (Jandt and Sigusch, 2009). While these have delivered some improvements in performance, one group has added SWCNTs to dental resin composite, citing their high tensile strength and Young’s modulus as the stimulus for their use. The work reported a process to deposit a thin shell of nano-silica onto the oxidised surface of SWCNTs by using a thin adlayer of 3-aminopropyltriethoxysilane, followed by modification of the layer by use of another organosilane with allyl-terminated functional groups with the aim of achieving good dispersion and integration of the nanotubes in the polymer matrix. The resultant dental resin composites gave significant increases in flexural strength (Zhang et al., 2008).
The aim of this study was to help disperse the SWCNTs within the matrix, by using the silica coating as a method to prevent direct tube–tube contact, reducing agglomeration and also improving their retention in the matrix under loading, by providing a rough surface. The authors reported that TEM showed that the thickness of the coating was around 10 nm and that the structure of the allyltriethoxysilane layer provided a compatible surface for further combination with the resin matrix. They also reported that the functionalised SWCNTs appeared as single strings between the filler particles of the dental resin composite and the matrix, implying that the composites can absorb more stress under applied load (Zhang et al., 2008). However, the resultant dental nanocomposites are described as being grey-black in appearance, which is obviously a major disadvantage for a dental restorative material.
The formation of the hard tissues of teeth depends on a natural process of biomineralisation. Methods have been described where clusters of needle-shaped apatite crystallites were grown on aggregated MWCNTs. The crystallites were 100 nm in width and 200–500 nm in length and were grown perpendicularly to the longitudinal axis of the nanotube and also radially, originating from a common centre of a single MWCNT (Akasaka et al., 2006). The apatite crystals were grown by immersing commercially obtained, and then purified, MWCNTs of curled shape in calcium phosphate solutions at a concentration of 10 mg/l. Initially the solution underwent ultrasonication for 10 minutes, and the apatite crystallites were then grown by immersion at 37 °C for various periods of time, up to 2 weeks. The authors stated that the MWCNTs may be acting as a core for the initial crystallisation of the apatites and that for nucleation to occur, an activation energy barrier must be exceeded by increasing the degree of supersaturation of the calcium phosphate solutions.
Periodontal dentistry is a dental specialty concerned with maintaining the health of the gingival tissues (gums) that support and attach the teeth to the jaw bones. In periodontal disease, chronic inflammation of the gingival tissues, due primarily to the presence of dental plaque containing millions of bacteria, occurs, which may lead to destruction of bone supporting the teeth, causing the teeth to become mobile, and eventually require extraction.
Kong et al. (2006) reviewed the potential use of different types of nanoparticles which may have practical applications in the treatment of periodontal disease. They speculated on the future use of CNTs for the local delivery of drugs with activity to slow or reverse the inflammatory process associated with periodontal disease together with the use of synthetic self-assembling scaffolds to replace lost bone tissue supporting the teeth. However, the use of CNTs in the possible treatment of periodontal disease has been more recently described by Mei et al. (2007) and Yang et al. (2009).
Mei et al. (2007) described the development of a new type of MWCNT containing composite material for use in guided tissue regeneration, where a biocompatible surface is inserted between the tooth root surface and the gingival tissue, which has lost its supporting bone through periodontal disease, with the ultimate aim of stimulating the re-formation of bone to support the teeth. The new material was a membrane created by electrospinning a suspension of poly(L-lactic acid), MWCNTs and hydroxyapatite. They found that the new membrane enhanced the adhesion and proliferation of periodontal ligament cells by 30% and inhibited the adhesion and proliferation of gingival epiteleial cells by 30% also (Mei et al., 2007). Clinically this may encourage the desirable regeneration and attachment of the tooth to bone, while discouraging the less desirable proliferation of gingival tissue which can interfere with reattachment of ligament cells.
Similarly, Yang et al. (2009) developed a new chitosan–MWCNT composite for potential use in guided tissue regeneration for the treatment of periodontal disease. MWCNTs have been shown to promote the adhesion and proliferation of osteoblasts but inhibit fibroblasts. Also, hydroxyapatite nucleates on oxidised MWCNTs functionalised with carboxyl groups. Yang et al. (2009) demonstrated the ability of apatite crystals to form on chitosan–MWNT composite materials. Figure 22.7 (p. 698) shows scanning electron micrographs of apatite crystals adherent to the surface of the chitosan–MWCNT composite material. The authors speculated that calcium ions bonded to negatively charged groups such as the carboxyl groups of MWCNTs, and more weakly with the amino groups of the chitosan, while phosphate ions associated with positively charged calcium ions. Apatite formation was deduced from characteristic XRD reflection peaks. They concluded that the method led to the formation of orientated nanoscopic crystallites of apatite which they assumed would enhance the bioactivity of the new material, presumably for use in guided tissue regeneration (Yang et al., 2009).
22.7 SEM images of apatite-treated chitosan-multiwalled CNT composite at various numbers of treatment cycles: (a) one cycle; (b) three cycles; (c) five cycles; (d) nine cycles and (e) an overall morphology of a failure surface following nine treatment cycles (Yang et al., 2009).
Dentures made to replace missing teeth in a dental arch are usually fabricated from PMMA. While exhibiting many advantages, such as excellent aesthetics, low density, and an ability to be repaired, it suffers from a relatively low fracture strength which renders it vulnerable to fracture, either in the mouth due to fatigue flexing, or due to accidental damage if dropped against a hard surface, such as a sink when being cleaned (Hargreaves, 1969; Jagger et al., 1999). Denture repair costs represent a significant cost to patients and health services. Many attempts have been made to increase the strength of PMMA by the addition of carbon fibres, glass fibres, metal plates, wires or mesh or chemical modification of the PMMA resin by addition of rubber graft copolymers (Vallittu, 1995; Vallittu, 1996; Jagger et al., 1999).
Recently work has been completed, incorporating MWCNTs into a PMMA-based denture base resin to determine if they could increase the strength and fracture toughness (Zhou, 2009). Unfunctionalised MWCNTs, carboxyl functionalised and amine functionalised MWCNTs at varied wt% loadings (0.1, 0.25, 0.5, and 1.0) were incorporated into the liquid monomer (prior to mixing) using ultrasonic agitation, and to the acrylic powder using dry blending and ultrasonic agitation, and the resultant specimens were heat cured. Mechanical characterisation was carried out including bending modulus, bending strength, compressive strength. Plane strain fracture toughness was also determined using the Chevron-Notch Short Rod method.
The authors reported that low loadings of MWCNTs (≤ 0.25 wt%) in PMMA denture resin significantly improved the mechanical properties of the resultant nanocomposite, while higher loadings (≥ 0.5 wt%) gave only small improvements, and in some cases significant reductions were recorded. The dispersion of MWCNTs within the matrix was enhanced by adding chemical functional groups, with carboxyl–MWCNTs giving the largest improvements. It was speculated that the improvements in mechanical properties were due to the MWCNTs being well dispersed within the matrix, thereby arresting/retarding crack propagation. Although the improvements in strength were significant, a disadvantage is the resultant black colour of the denture base resin, which would require masking if it were to be used clinically.
Squamous cell carcinoma is one of the more common types of mouth cancer and is responsible for an estimated 650,000 new cancers annually (Argiris et al., 2008). Treatments that selectively target the cancer cells are required as current treatments lack specificity and can cause severe side effects. Recently oxidised SWCNTs have been bioconjugated with the anti-cancer drug cisplatin and a specific receptor ligand, epidermal growth factor (EGF) (Bhirde et al., 2009). EGF has a strong affinity for a cell-surface receptor which is over-expressed in most squamous cancer cells. The SWCNTs are therefore targeted at the cancer cells improving specificity of treatment. Microscopy showed that the functionalised bioconjugated nanotubes interacted with the cell surface receptors which in turn caused endocytosis, drawing the nanotube into the cell, leading to cell death. When tested in mice, administration of the SWCNT–cisplatin–EGF bioconjugate caused a very significant reduction in tumour volume, compared to SWCNT–cisplatin control groups (Bhirde et al., 2009).
Carbon nanotubes are being increasingly used in the fields of regenerative medicine and tissue engineering. As human stem cells are introduced into the body, there is a need to be able to accurately label them and track their movement as they divide, differentiate and migrate. CNTs are proving useful in this field. In addition, CNTs are also being investigated to monitor cellular behaviour, augmenting cellular behaviour (drug delivery) and enhancing the mechanical properties of tissue scaffolds used to support newly introduced cells (Harrison and Atala, 2007).
It initially proved difficult to track the movement of CNTs in biological systems using conventional methods such as by elemental analysis, as they only contain carbon, and also by electron microscopy due to their small size and relatively low contrast, so alternative methods were sought to overcome these disadvantages.
SWCNTs have been covalently linked to visible-wavelength fluorophores and imaged within cells (Pantarotto et al., 2004; Kam et al., 2004). However, Cherukuri et al. (2004) highlighted some potential disadvantages of this system. These were that the chemical linkage must be able to resist enzymatic cleavage, emission from the visible wavelength fluorophore must be detected above background endogenous fluorescence, and that chemical processing of the CNTs may alter their biological fate. That is, that the process of labelling the CNTs for tracking may significantly alter their behaviour in the cell and give erroneous results. To overcome these problems they proposed a method for observing pristine SWCNTs in biological systems using their intrinsic near-infrared (NIR) fluorescence. SWCNTs were incubated with cultured mouse peritoneal macrophage-like cells, in a growth medium. The incubated cells were analysed for NIR fluorescence using a spectrofluorometer and a fluorescence microscope. Fluorescence was excited by light from a 660 nm diode laser and showed distinct emission peaks arising from the semi-conducting SWCNTs. CNTs possess many desirable properties for optical detection in biological systems. They display optical transitions in the NIR spectrum between 900 and 1300 nm, which is an important optical range for biomedical applications because of the good penetration depth of light and small auto-fluorescent background, together with excellent photostability (Bruchez, et al., 1998; Chan and Nie, 1998). Heller et al. (2009), similarly used SWCNTs as optical sensors to detect genotoxic analytes, including chemotherapeutic drugs in real time within live cells.
CNTs are useful for labelling cells as their small size means that they are less likely to cause physical damage to the cells under investigation, potentially altering their subsequent behaviour (Harrison and Atala, 2007). Chen et al. (2007) described the development of a single MWCNT attached to an atomic force microscope tip and functionalised with quantum dots (light-emitting nanocrystals composed of atoms from groups II–VI of the Periodic Table), by a disulphidebased linker. Penetration of cell membranes with the MWCNT was controlled by an atomic force microscope, and single-particle tracking was used to trace the movement of the quantum dots within the cell.
Tissue engineered constructs may also be monitored over time using magnetic resonance imaging (MRI). Unmodified CNTs do not contrast well against body tissue in an MRI scan, however they can be functionalised to make them more visible, for example by the addition of gadolinium heavy atoms (Hartman and Wilson, 2007; Richard et al., 2008). Sitharaman et al. (2005) showed that nanotubes 20–100 nm long could be loaded with gadolinium through the openings at the end of the nanotubes, or defects in the walls, providing high contrast in MRI scans, enabling monitoring of the engineered tissue. Choi et al. (2007) described the formation of hetero-structured complexes of magnetic iron oxide nanoparticles and NIR fluorescent SWCNTs that could be used as multimodal bio-imaging agents. The resulting nanotube complexes showed near infra-red (NIR) fluorescence, Raman scattering and visible/NIR absorbance features. Macrophage cells that engulfed the nano-complexes were imaged using MRI and NIR mapping, showing that the multifunctional nanostructures could be potentially useful in bio-medical imaging.
Radiotracers have been added to CNTs to enable them to be tracked using gamma scintigraphy (Singh et al., 2006). This study followed the bio-distribution of nanotubes in mice and found that functionalised SWCNTs did not accumulate in any particular organ, and were rapidly cleared in the urine.
Tissue engineering requires not only knowledge regarding the distribution of implanted stem cells, but also of the environment and reactions occurring within the cells. The ability to monitor cellular physiology including ion transport, enzyme/cofactor interactions, protein and metabolite secretion could be used to monitor the performance of cells within engineered and normal tissues. CNTs are ideal components for nanosensors due to their small size, electrical properties and large surface area for binding to many biological compounds, including DNA and proteins (Harrison and Atala, 2007). CNT-based sensors have been used to measure the electro-oxidation process of insulin using MWCNTs. Other substances detected by CNT based sensors include amino-acids, putrescine, cholesterol, together with physiological conditions such as pH (Radosavljevic et al., 2002; Xian et al., 2005; Rochette et al., 2005; Tan et al., 2005).
The above methods rely largely on electrochemical sensing to detect substances, but CNT-based sensors may also exploit changes in optical properties, for example, to measure the levels of β-D-glucose concentration (Barone et al., 2005) by a change in the near infra-red fluorescence of the SWCNT sensor.
Kim et al. (2007) reviewed the use of CNTs for the electronic and electrochemical detection of biomolecules such as enzymes, proteins and DNA, together with strategies for amplifying the resultant signals to enable sensitive detection down to picomolar range. The intrinsic bandgap in the density of states of semiconducting SWCNTs allows them to act as semiconducting nanosized channels in a field-effect transistor (FET) (Chen et al., 2003; Hartman and Choi, 2006; Star et al., 2006). The SWCNTs detect changes in their environment due to the specific interactions with the biomolecules of interest. FETs can be further divided into three types, dependent on the receptor type and how a signal is generated:
3. Immunologically modified FETs and DNA-modified FETs which use surface polarisation effects or changes in the dipole moment, e.g. antigen–antibody binding or DNA hybridisation (Uno et al., 2007).
22.8 Structure of an FET nanobiosensor. (a) Cross-sectional view: source and drain electrodes bridge the semiconductor channel. The gate electrode can be used to modulate the conductivity of the semiconductor channel. A receptor molecule attached to the surface of the semiconductor material can specifically recognise and capture a target molecule from a buffer solution. (b) Top view: SEM image of a typical nano-FET. In these structures, the channel length is the S–D distance and the channel width is the S or D electrode width. Examples are shown of nano-FET fabricated at the author’s research facility using either (c) carbon nanotubes or (d) indium oxide nanowires as semiconductor materials (Curreli et al., 2008).
CNTs have also been used as electrodes in electrochemical-based nanobiosensors, boosting direct electron transfer to aid detection of important enzymes such as glucose oxidase, which have previously been difficult to measure using conventional methods. Kim et al. (2007) described the two approaches that have been used. In the first method, SWCNTs or MWCNTs are randomly deposited onto conductive surfaces in a mat configuration, or packed into a micropipette for use as electrodes. The second method involves the creation of SWCNT forests, with shortened SWCNTs standing vertically, with one end in contact with the electrode and the other exposed to the test solution (Kim et al., 2007).
In conclusion, CNTs have been demonstrated to be of great use as biosensors, despite exhibiting some disadvantages such as the need to separate semiconducting nanotubes from metallic nanotubes and a non-uniform distribution of bandgaps which may lead to difficulty in fine-tuning electronic properties (Curreli et al., 2008). The use of CNTs as biosensors has been well reviewed in a number of papers (Balasubramanian and Burghard, 2006; Gruner, 2006; Allen et al., 2007; Kim et al., 2007; Curreli et al., 2008).
CNT polymer composites used for biomedical and bioengineering applications require special consideration regarding their impact on the human body. Information is required regarding the route by which CNTs and their composites are introduced into the body and how they subsequently move through the body in the short and long term.
Much of the work done to assess the biocompatibility or toxicity of CNTs used pristine, unfunctionalised CNTs that were dispersed in different media and then introduced into animal models by inhalation into the lungs or injection. The results of these tests have been very mixed with some indicating significant risks associated with specific types of CNTs and others not. However, this confusion is understandable when the large diversity of types of CNT tested is considered. Whether single-walled or multi-walled, functionalised or pristine, free within solution or as part of a polymer composite, coated or uncoated, agglomerated or dispersed, short or long, straight or tangled, they will behave differently in different experimental set-ups (Helland et al., 2007). Currently there are no internationally agreed protocols for testing the biocompatibility of CNTs and several authors have highlighted the urgent need for them.
CNTs have attracted much interest among those concerned regarding the potential toxicity of nanotubes due to the physical similarity of some types of CNT fibre to asbestos fibres. The harmful effects of exposure to brown asbestos fibres are well known in the pathogenesis (development) of a type of lung cancer, a mesothelioma (Mossman and Churg 1998).
In 2005, Murr et al. found that exposure to SWCNTs or MWCNTs caused asbestos-like cytotoxicity determined using the MTT assay. The MTT assay is a quantitative colorimetric method to determine cell proliferation and viability. It uses a yellow tetrazolium salt [3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium-bromide] which is metabolized by mitochondrial succinic dehydrogenase enzyme activity from proliferating cells, producing a purple formazan reaction product. Similar results have been reported by other workers (Soto et al., 2005; Jia et al., 2005; Soto et al., 2007).
Previous work on asbestos fibres had shown that any fibre is likely to be harmful to human lungs if it is thin enough to enter the lungs (< 3 microns), longer than the lung cells that usually remove foreign fibres (> 20 microns) and is insoluble in the lungs, and therefore likely to accumulate (Donaldson and Tran, 2004). Poland et al. (2008) investigated if this paradigm also applied to CNTs and found that exposing the mesothelial lining of the chest cavity in mice to long MWCNTs resulted in asbestos-like adverse effects. These included inflammation and the formation of granulomas (Poland et al., 2008). The paper emphasised that, although their data showed that short CNTs did not mimic the behaviour of long asbestos, the possibility that short CNTs are also potentially harmful to humans by some other mechanism, such as intrinsic toxicity, could not be excluded. In addition, the authors stated that it is unknown whether the inflammatory and granulomatous changes in the tissues induced by the long CNTs would go on to develop a mesothelioma tumour. In addition, the degree of metal catalyst contamination of the CNTs tested has been found to be a factor in the magnitude of in vitro cytotoxicity displayed. Samples with higher levels of iron contamination (30%) or nickel (20%) caused significant decreases in cell viability (Kagan et al., 2006; Herzog et al., 2007; Shvedova et al., 2010). Conversely, CNT samples containing low levels of iron were less toxic to cells (Kagan et al., 2006; Herzog et al., 2007). This increased cytotoxic effect may be because MWCNTs have been shown to induce cellular oxidative stress, which was removed when the CNTs were acidtreated removing the metal contaminants (Pulskamp et al., 2007).
In contrast to these results, other studies have confusingly reported low cytotoxicity for raw SWCNTs with high iron contamination, purified SWCNTs with low iron content and purified MWCNTs (Worle-Knirsch et al., 2006; Pulskam et al., 2007).
The potential toxicity of MWCNTs may also be influenced by the presence of structural defects in their carbon framework (Fenoglio et al., 2008). Shvedova et al. concluded in their review paper that many physico-chemical characteristics of MWCNTs, e.g. length, width, structural defects, metal contamination, surface chemistry and zeta potential will affect the lung response. Further, they stated that until the exact parameters which contribute to their toxicity are characterised, different types of MWCNT materials will give differing biological responses (Shvedova et al., 2009).
Shvedova et al. (2009) have suggested that these conflicting results may be more comprehensible by considering that the dye-based cell viability assays used may be interfered with by the CNTs due to light absorption and scattering. Therefore they suggest that results of cell viability studies, based on MTT dye-based assays, should be interpreted with caution. Recent studies have investigated SWCNT uptake by different cells. Research shows that CNTs are able to cross the cell membranes of rats and therefore might have an influence on cell function (Pulskamp et al., 2007). The results obtained have varied and it appears that the presence or absence of specialised signals determines the recognition and subsequent interactions of SWCNTs with cells. Overall, pristine SWCNTs carrying no recognisable signals are poorly taken up whereas those modified chemically e.g. oxidatively modified during purification and storage, or by adsorbed macromolecules such as proteins, are more readily recognised and engulfed by cells (Shvedova et al., 2009). Several studies have shown that SWCNTs are not readily taken up by lung cells (Shvedova et al., 2005; Herzog et al., 2007; Davoren et al., 2007), while others have reported uptake of both SWCNTs and MWCNTs by rat and guinea pig cells (Jia et al., 2005; Dutta et al., 2007; Pulskamp et al., 2007).
Dutta et al. (2007) reported that CNTs highly adsorb albumin when placed in a serum dispersion medium. This process structurally alters the albumin which triggers scavenger receptors. Therefore the reported uptake of CNTs may be due to the altered albumin coating rather than the CNT alone. The choice of dispersion medium for CNTs is therefore important (Porter et al., 2008).
The potential toxicity of CNTs to the human immune system also requires consideration. Research has shown that nanoparticles can stimulate and/or suppress the immune system, and that the response is largely determined by their surface chemistry (Mitchell et al., 2007; McDonald and Mitchell, 2008). Overstimulation of the human immune system may lead to inappropriate allergic-type reactions, while suppression may adversely alter the normal immune response to infections and cancer cells.
Exposure to SWCNTs by inhalation may also interfere with the body’s ability to eliminate infection. Shvedova et al. found that exposure to CNTs and infection by Listeria monocytogenes induced unusual responses in mice with mutually enhanced inflammation, together with depressed bacterial clearance (Shvedova et al., 2008). Therefore exposure to CNTs through inhalation may exacerbate lung infections in susceptible populations (Shvedova et al., 2009).
Concerns exist that materials containing CNTs could alter the genetic material of humans and other organisms, leading to the potential of carcinogenic changes. Research has shown that both SWCNTs and MWCNTs may cause changes to the genetic material but the results obtained to date are not consistent (Shvedova et al., 2009).
The movement of nanoparticles within the body after initial introduction by inhalation, intravenous injection, or ingestion, has been investigated for various nanoparticles, such as ultrafine carbon particles, and found that labelled particles were found in various organs throughout the body of rats. One study specifically investigated the translocation or movement of MWCNTs within mice. Within 10 minutes of intravenous injection of 14C-taurine-labelled MWCNTs were found mainly accumulated in the liver and also the heart and lung, but not the brain, stomach, muscle, bone or intestines (Deng et al., 2007).
The potential for humans and the environment to be exposed to CNTs during manufacture, distribution and use/disposal of products and devices containing both SWCNTs and MWCNTs is of intense interest but little is known regarding this (Maynard and Kuempel, 2005). Air-borne levels of MWCNTs during manufacturing ranged from undetectable to 400 g/m3 (Han et al., 2008). Spraying, blending or weighing produced the highest concentrations, but the implementation of controls significantly reduced levels to often non-detectable levels. Maynard et al. (2004) reported a field study in which airborne and dermal exposure to SWCNTs was investigated while handling unrefined material. Although laboratory studies indicated that with sufficient agitation, unrefined SWC NT material could release fine particles into the air, concentrations generated while handling material were very low with estimates of airborne concentration of nanotubes lower than 53 μg/m3 in all cases (Maynard et al., 2004).
Several nanotoxicology workshops and symposia have been held to discuss the potential problem of adverse environmental and human effects of CNTs (University of Florida, 2004; Tsuji et al., 2006; Balbus et al., 2007). In general it is agreed that health and safety and environmental protection laws internationally lag behind the manufacture and use of CNT-containing products, devices and materials. Progress in this area has been delayed by discussion regarding whether nanoparticles with the same chemical composition of bulk or macro counterparts, such as graphite in the case of CNTs, were new materials requiring new legislation, or existing materials with a significant new use.
Research has shown that nanoparticles are more hazardous than their bulk counterparts. In addition to particle size, surface area various physico-chemical factors including surface reactivity, solubility and shape can influence the toxicity of engineering nanoparticles (Shvedova et al., 2009).
The need to proactively adopt appropriate risk management procedures for nanotubes, including CNTs, has recently been reviewed (Murashov and Howard, 2009). They highlighted the need for an adaptive approach to risk management that should exist in ‘real time’, by used of web-based platforms for the development of consensus-based dynamic global standards, e.g. the wiki-based project ‘GoodNanoGuide’ (GoodNanoGuide, 2009). Many different international organisations are developing guidelines for the safe handling of nanomaterials and the collection of risk information (Nanosafe, Institute of Occupational Medicine, SAFENANO, Institute for Work and Health, National Institute for Occupational Safety and Health, Internano, The Organisation for Economic Co-Operation and Development).
The requirement for new technologies and medical implants in the treatment of traumatic injuries and chronic diseases is increasing. This is due in part to an ageing population with the related degenerative processes and disease processes, and also to changing daily lifestyles with a rising need for solutions to healthrelated issues. Novel concept biomaterials and medical devices are urgently needed to allow less tissue damage and more tissue regeneration, likewise such materials and devices need to be implanted under minimally invasive conditions, conducive to a rapid recovery. Particularly for biomaterials and devices designed to replace a degenerated or diseased joint, bone or tooth structure, many questions need to be answered. Such devices and implants would benefit significantly from availability of a material that is multi-functional and can meet the biomechanical and biological requirements.
The conventional biomaterials available today are reaching their maximum capabilities, notwithstanding their successful application in treating and preventing different medical conditions. There is a need for the development of new biomaterials which must satisfy several requirements ranging from physical, mechanical, biological, toxicological and other characteristics, depending on the final clinical application.
Carbon is chemically inert and CNTs not only demonstrate superior mechanical, chemical and electrical properties, but also have the potential to be biocompatible particularly when functionalised. Also, encapsulation of other materials within CNTs could potentially create applications for therapeutic use in medicine.
Regardless of this interest, there are many factors and limitations to be borne in mind. The field of nanomaterials for biomedical and bioengineering applications is still very much in its early stages and many difficult questions must be addressed, including manufacturing, safety and regulatory issues. Preliminary investigations substantiate the enormous potential of CNTs for biomedical and bioengineering applications either as a structure, coating, scaffold or composite, although most of these are only at laboratory-scale and in vitro testing. There is a major requirement for interdisciplinary collaboration and exchange of knowledge at many levels to effectively address the current issues, before being able to fully understand and explore the true potential of CNTs for biomedical and bioengineering applications.
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