In 2012 alone, among the total of 86,488 hip replacement procedures in England, Wales and Northern Ireland, over 10,000 were revision procedures [6]. This revision procedure accounts to 12% of total hip replacement in 2012 whereas in 2011 it was 11%. There are many reasons for the failure of an implant within the biological environment including manufacturing, chemical, mechanical, tribological and surgical failures [3]. The patient’s health
condition and the physician’s experience could also be contributing factors. A
variety of biomaterials (polymers, metals, ceramics and composites), fabrication techniques (such as compression moulding, die casting, bar stock milling and laser cutting) and surface modification techniques (biocompatible material coating, surface polishing and drug coating) have been employed to improve the mechanical and biological properties of the implant; however, there are still constraints in achieving this objective.
Chapter 1: Introduction
Even when using a biomaterial with suitable material and biological properties, there are possibilities for the failure of an implant due to faulty mechanical design or inappropriate application of the implant. Inadequate mechanical properties (e.g. elastic modulus, yield strength, tensile strength) can result in fracture leading to implant failure [17,28,29]. Elasticity modulus is the resistance offered by an object/substance when deformed elastically. Yield strength is the stress at which a material starts to deform plastically. Tensile strength is the maximum stress which the material can withstand before its failure. Fracture of an implant mainly occurs when the implant is not capable of bearing the load exerted on it. For example, the major purpose of orthopaedic implants is to restore the function of load bearing joints (such as hip or knee joints) that are subjected to high levels of mechanical stresses, wear and fatigue during the course of normal activity [4]. According to Wolff’s law, the bone of a healthy human or an animal will adapt to the loads under which it is placed. For example, when there is an increase in loading of a particular bone, the bone will remodel itself over time to resist that load. The internal structure of the trabeculae will undergo adaptive changes, leading to secondary changes to the cortical portion of the bone. However, it should be noted that if the loading on a particular bone decreases, then the bone will become weak.
Although the mechanical properties of the bulk material, implant design and manufacturing process are considered to be the major factors determining implant fracture, corrosion of implants within the biological environment is also a crucial factor [3]. Before the advent of 316 stainless steel, metallic implants of various steel formulations failed dramatically due to tissue reactions [30]. 316L SS have a better corrosion resistance and relatively low carbon concentration in the alloy mixture compared to other steels. The presence of molybdenum increases the corrosion resistance of the alloy.
Implants corrode within the biological environment due to electrochemical attack by the electrolytes present in the hostile environment. Body pH can affect the corrosion resistance of implants. Although the pH value of the human body is normally maintained at pH 7, this might change from pH 3 – 9
Chapter 1: Introduction
depending on the imbalances in the biological system due to infections, diseases and other factors. The aqueous medium in the human body consist of anions including chloride, phosphate and bicarbonates, and cations including sodium, potassium, calcium and magnesium [3]. Leaching of metal ions from the implant surface due to corrosion will lead to erosion. The rate of corrosion is accelerated by increased surface area and loss of protective oxide film, later leading to fractures. Also, leaching of metal ions has been observed to induce acute and chronic effects. For example, the release of nickel from stainless steel was observed to affect the skin; cobalt was observed to cause anaemia which inhibits the absorption of iron into the blood stream; chromium was observed to cause ulcers and disturb the central nervous system; aluminium
was reported to cause epileptic effects and Alzheimer’s disease and leaching of
vanadium in its elemental state was observed to be toxic [31]. The tolerable corrosion rate for metallic implants is 2.5 x 10-4 mm/year [3]. Hence, corrosion products formed as a result of implant-host interaction have a significant effect on the long term stability of the prosthesis and its cytocompatibility [32].
Biocompatibility, a requirement for all biomedical implants, is the property of a material to be compatible with the tissues in the human body by not producing a toxic, injurious or immunological response [4]. Chronic inflammations, lesions and scarring are some of the adverse reactions that may take place in the site if the implant is not biocompatible. To achieve better biocompatibility, two routes are often employed: (i) using biocompatible materials such as gold; (ii) fabrication of implants using off-the-shelf materials with application of a suitable biocompatible layer such as bio-active glass, calcium phosphate and protein coatings [15]. The corrosion resistance of titanium is better than other metals such as 316L SS due to the presence of a stable surface oxide layer (TiO2) of approximately 10 nm. However, the
thickness of this surface oxide layer can be enhanced by various surface modification methods including chemical etching and anodisation [30].
Contamination is another factor contributing towards an implant failure. Contamination of implants such as stents, dental, hip and knee implants by harmful microbes has been widely reported [33–36]. Bacterial infections due to
Chapter 1: Introduction
implants are one of the main risk factors in orthopaedic surgery leading to surgical removal of implants [17,37,38]. Contamination of an implant is possible due to non-sterile manufacturing and packing, during surgical placement and/or even due to the microbes present in the body after implantation into the body [38]. By entering the host through contaminated surfaces, bacteria multiply in the host environment and interfere with the host defence system leading to host tissue damage and inflammation [38]. Staphylococci species, including Staphylococcus aureus and Staphylococcus epidermidis are the more common cause of implant-associated infections [39].
Treatment measures available for bacterial infections include systemic, antibiotic loaded cements and antimicrobial coatings [37]. However, the release of drugs in a systemic therapy prevails as one of the major problems when conventional coating techniques are used. Slower release of antibiotics for a longer period of time is a potential cause of antibiotic resistance in the human body [40]. For example, in drug eluting stents (DES), to reduce the restenosis rate, the stents are coated with polymers containing a drug using simple techniques including dip/spray/spin coating or solvent casting [41,42]. After implantation, the DES is expanded to conform to the vessel wall. During this expansion, fractures, peeling of the polymeric coating and uneven release of the drug material were reported as some of the major problems with surface modified DES [43,44]. To address these issues, Mani et al. [45] adopted a coating technique by which the drug was directly loaded on to the implant surface. However, the amount of drug coated was very low (~ 5µg/cm2) compared to the amount of drug coated on a commercially available drug- eluting systems (100 µg/cm2). Polymer brushes were also used to coat implant surfaces with drugs [46]. Since these coating methods are in nano-scale, they add only a few nanometres to the implant surface and hence, peeling of drug coating is less likely to occur [47]. Also, implants coated with conventional coating techniques have drawbacks including surface heterogeneity in the type and distribution of functional groups, hydrophilic and hydrophobic domains and surface roughness [48].
Chapter 1: Introduction