Name Greg. Ulonna
Degree or QualificationMSc
Thesis/Dissertation/AssignmentBIOMATERIALS ASSIGNMENT-Uni. of Wolverhampton-U.K (not well researched) (1) Introduction 1 Nowadays, there are several bio-implant materials available for use in total joint replacements (TJRs) as well as for knee and hips replacements. Of these, metallic alloys like austenitic stainless steels and Co–Cr–Mo alloys [1] [1] are preferred for load bearing components. In the United Kingdom, the most popular alloy is Ortron 90, an austenitic Fe–20Cr–10Ni–2.5Mo–0.4N steel, which is used for both one-piece femoral components or in modular form as a femoral head. The former is a well-proven and highly reliable device that was originally designed by Sir John Charnley [2] [1]. It remains the most popular of those used by surgeons in the UK. However, in the USA, and elsewhere in the world, Co–Cr–Mo alloys are favoured, especially for modular femoral heads. In principle, the titanium alloy Ti–6Al–4V, is also very attractive, especially for one-piece femoral components. It has a high specific strength and a well-known “bone-friendly” character. [1]. Other materials include a nanostructured hydroxyapatite (HA) feedstock thermally sprayed on Ti–6Al–4V substrates via high velocity oxy-fuel (HVOF process), as well as polymers. Today, nearly one out of ten Americans use plastic contact lenses to correct vision. [2]. Everyday, roughly 45,000 Americans with kidney failures get hooked up to artificial kidneys, wherein polymeric membranes remove life-threatening toxins from blood. [2]. Some 7,000 angiography procedures are performed in the U.S. each day, to determine patients' cardiovascular health. [2]. A thin plastic tube, known as a catheter, is snaked in a patient's blood vessel, and radio-opaque dye is injected. [2]. When blood flow is obstructed, catheters equipped with tiny inflatable balloons are usually used to dislodge the plaque deposits obstructing the arteries. [2]. Such modern medical plumbing often prevents expensive and more invasive surgery. [2]. When major surgery is called for, surgeons perform bypass surgeries; often using synthetic plastic blood vessels to bypass damaged large arteries. [2]. When bypassing damaged coronary arteries, more often than not, it is a membrane oxygenator, with a polymeric membrane, that takes over the function of lungs during these open-heart procedures. [2]. Materials, like the ones mentioned previously and particularly synthetic polymers, more than anything else, have been instrumental in improving the quality of health-care the world over. (2) Literature Review Use of biomaterials dates far back into ancient civilizations. Artificial eyes, ears, teeth, and noses were found on Egyptian mummies [256][3]. Chinese and Indians used waxes, glues, and tissues in reconstructing missing or defective parts of the body. Over the centuries, advancements in synthetic materials, surgical techniques, and sterilization methods have permitted the use of biomaterials in many ways [178][3]. Medical practice today utilizes a large number of devices and implants. Biomaterials in the form of implants (sutures, bone plates, joint replacements, ligaments, vascular grafts, heart valves, intraocular lenses, dental implants, etc.) and medical devices (pacemakers, biosensors, artificial hearts, blood tubes, etc.) are widely used to replace and/or restore the function of traumatized or degenerated tissues or organs, to assist in healing, to improve function, to correct abnormalities, and thus improve the quality of life of the patients. [3]. According to a report published in 1995 by The Institute of Materials, London, the estimated world market for all medical devices, including diagnostic and therapeutic equipment is in the 2 region of $100 billion per year. [3]. Within this industry, the world market for biomaterials is estimated to be around $12 billion per year, with an average global growth of between 7 and 12% per annum. [3]. Where Biomaterials are expected to perform in our body's internal environment, are very aggressive environments. For example the pH of body fluids in various tissues varies in the range from 1 to 9. [3]. During daily activities bones are subjected to a stress of approximately 4 MPa whereas the tendons and ligaments experience peak stresses in the range 40–80 MPa. [3]. The mean load on a hip joint is up to 3 times body weight (3000 N) and peak load during jumping can be as high as 10 times body weight. [3]. More importantly, these stresses are repetitive and fluctuating depending on the activities such as standing, sitting, jogging, stretching, and climbing [21][3]. In a year, the stress cycles of finger joint motion or hip joint motion estimated to be as high as 1×106 cycles, and for a typical heart 0.5 ×107–4×107 cycles. This information roughly indicates the acute and instantaneous biological environment in which the biomaterials need to survive. Needless to say, the biological environment also depends on the patient's conditions and activities. [3]. In the early days all kinds of natural materials such as wood, glue and rubber, and tissues from living forms, and manufactured materials such as iron, gold, zinc and glass were used as biomaterials based on trial and error. The host responses to these materials were extremely varied. The body tolerated some materials whereas others were not, but certain conditions (characteristics of the host tissues and surgical procedure) some materials were tolerated by the body, whereas the same materials were rejected in another situation. Over the last 30 years considerable progress has been made in understanding the interactions between the tissues and the materials. It has been acknowledged that there are profound differences between non-living (avital) and living (vital) materials. [3]. Researchers have coined the words ‘biomaterial’ and ‘biocompatibility’ [253][3] to indicate the biological performance of materials. Materials that are biocompatible are called biomaterials, and the biocompatibility is a descriptive term, which indicates the ability of a material to perform with an appropriate host response, in a specific application [22][3]. In simple terms it implies compatibility or harmony of the biomaterial with the living systems [258][3], extended this definition distinguishes between surface and structural compatibility of an implant [260] [3]. Surface compatibility meaning the chemical, biological, and physical (including surface morphology) suitability of an implant surface to the host tissues. Structural compatibility is the optimal adaptation to the mechanical behaviour of the host tissues. Therefore, structural compatibility refers to the mechanical properties of the implant material, such as elastic modulus (or E, Young's modulus) and strength; implant design (stiffness, which is a product of elastic modulus, E and second moment of area, I), and optimal load transmission (minimum interfacial strain mismatch) at the implant/tissue interface. [3] Optimal interaction between biomaterial and host is reached when both the surface and structural compatibilities are met. Further more it should be noted that the success of a biomaterial in the body also depends on many other factors such as surgical technique (degree of trauma imposed during implantation, sterilization methods, etc), health condition and activities of the patient. [3]. (3) Critical Analysis 3 Biomechanical/Biomedical analysis uses the theories and methods of mechanics and mechanical engineering in conjunction with anatomical and physiological knowledge to analyze injuries. [4]. Using the biomechanical techniques, engineers analyze an injury or a general form of injury or impairment whether involving a vehicular accident, recreational injury, or workplace-related injury or even as a result of a natural breakdown or change. Achieving a long-term stable implant interface is a significant clinical issue when there is insufficient cortical bone stabilisation at implant placement [5], a stability human study and endeavour cannot adequately offer. Clinical outcome studies suggest that high-risk exists especially for implants placed in compromised cortical bone (thin, porous, etc.) in anatomical sites with minimal existing trabecular bone (characterised as type IV bone). [5]. In establishing and maintaining an implant interface in such an environment, one needs to consider the impact of masticatory forces, the response of bone to these forces and the impact of age on the adaptive capacity of bone [5], and of course these impacts can never be sufficiently substantiated as the very and total organic make up of the human being is still not adequately established to begin with. These forces, in turn, have the potential to create localised changes in interfacial stiffness through Viscoelastic changes at the interface. [5]. Changes in bone as a function of age (e.g. localised hypermineralised osteoporosis and localised areas of osteopenia) will alter the communication between osteocytes and osteoblasts creating the potential for differences in response of osteoblastic cells in the older population. [5]. Implants made of polymerics are usually susceptible to deteriorations from inherent or naturally occurring acids in the human system, and what follows afterwards can be damaging up to unprecedented scopes. The Thermal compatibility of Metal-Ceramics a material used for dental prostheses for instance has not been satisfactorily determined to predict if such a material system will be susceptible to failures in clinical situations. Such are the issue abound in the application of implants in Biomedicine and Bioengineering. (4) Possible failure modes Joints usually fail when applied stresses surpass fatigue limits. Trays like tibial trays are likely to fail if half of it is clamped and the other unsupported half carries the load, this would be a severe condition if it were to occur in a patient, requiring a reduction in physiological load. [6]. In real life, both sides of the tibial component provide support, which is why many artificial knee joints survive actual loads of the order of 2000 N [6]. Based on a test conducted by Sunita Ahir PhD for tibial trays, the survival of such implants can be compromised once fibrous tissues or arthritic conditions develop. Fibrous tissue can develop under the base plate, leading to increased bending of the tray caused by the more-compliant support in comparison with the hard bone [6]. According to her “This can lead to a fracture of the tibial component in poorly designed implants”. [7]. Titanium and its alloys, particularly Ti–6Al–4V, have seen significant service as load bearing bio-implant materials [1][7]. They are regarded as 4 being "bone friendly" and consequently achieve good osteointegration in many situations such as certain orthopaedic femoral stem components or dental implants where fixation cement is not required. [7]. However, there is always a concern regarding implant loosening. [7]. The causes are complex and, in the case of dental implants, are improperly understood. However, the creation of wear debris through micro-motion at the implant-bone interface or by sliding contact between, for example, a femoral head component and an acetabular cup, exists and is to be avoided. [7]. Based on conducted corrosion–wear tests reported by P. A. Dearnley et al; untreated CP-Ti and Ti–6Al–4V alloys principally degrades through micro-asperity shearing—a rapid process. Type I corrosion–wear, although possible, is not predominant. [7]. There was no evidence of Type II or Type III corrosion–wear processes after testing the oxidised CP-Ti and Ti–6Al–4V alloys in 0.89% NaCl. [7]. The processes responsible for the degradation of the exterior TiO2 (rutile) formed on the oxidised CP-Ti and Ti–6Al–4V alloys differed. [7]. In the latter case, rapid loss of the layer occurred during the first 60 min of testing. [7]. Failure took place via spallation, i.e., fracture along the TiO2–ODZ interface. [7]. This was attributed to high levels of residual stress in the TiO2 layer. [7]. In the case of the TiO2 layer formed on the CP-Ti alloy no spallation took place, since residual stress levels were lower. [7]. Instead, the TiO2 layer was slowly worn through. [7]. The track surface was smooth, suggesting degradation by a combination of abrasion and corrosion–wear processes. [7]. The latter, possibly involved the mechanical dissociation of the TiO2 layer, a previously unreported phenomenon. [7]. Once the TiO2 layers were removed from the oxidised Ti–6Al–4V and CP-Ti alloys, the adjacent hardened ODZ was smoothly worn, at a rate that was much lower than for the untreated Ti test-pieces. [7]. It is proposed that the principal degradation was caused through Type I corrosion–wear, whereby the passive film is repeatedly removed and reformed during sliding contact with the -Al2O3 ball slider. [7]. Two types of wear were noted for the -Al2O3 ball sliders: (i) a smooth wear process caused by micro-abrasion when in contact with the oxidised CP-Ti and Ti–6Al–4, and (ii) a rougher wear process resulting from grain pull-out, a phenomenon that was only observed when in contact with the untreated CP-Ti and Ti–6Al–4 test surfaces. [7]. Fig. 1. Type I corrosion–wear [10][7]. 5 Fatigue fracture and wear have been identified as some of the major problems associated with implant loosening, stress-shielding and ultimate implant failure [1][9]. Although wear is commonly reported in orthopaedic applications such as knee [2][9] and hip joint [3][9] prostheses, it is also a serious and often fatal experience in mechanical heart valves [4][9]. Fig. 2 illustrates some examples of fatigue fracture of implant devices in the hip prosthesis and a mechanical heart valve. It can be seen that fatigue-wear interaction plays a significant role in ultimate failure of these medical devices. [9]. The acetabular cup made of ultra-high molecular weight polyethylene (UHMWPE) has been worn so severely that it fractured in a brittle manner. [9]. This was in spite of its relatively high initial fracture toughness in the order of 2 MPa m. [9]. The cast cobalt chrome femoral stem has fatigue fractured at its lower proximity. [9]. The polyacetal occluder of the tilting disk heart valve shows a deep wear groove as a result of the repetitive impact-cum-sliding motion it made with the upper metallic strut. [9]. Fig. 2. Some examples of fatigue failure of medical devices: (a) hip prosthesis; (b) explanted Björk–Shiley polyacetal disc mechanical heart valve (arrow indicates fatigue-wear mark). [9]. (5) Sustainability of Implants According to a financial consulting firm, the worldwide market for electronic medical implants is estimated to reach $9.68 billion in 2004, and growth will continue at a compound rate of 14% to reach approximately $18.65 billion in 2009. The economic costs associated with device recalls totalled nearly $1 billion over the past decade. [8]. These costs include only hospitals stays, device replacement, and doctor’s fees; and not the cost of pain, suffering, and loss of life. [8]. In other to adequately determine what the sustainability of Implants is some life cycle analysis 6 has to be carried out; some study carried out show ultra-high molecular weight polyethylene (UHMWPE) and polyacetal are not exactly sustainable as they contribute to problems in wear debris formation and failure especially in acetabular UHMWPE cup used in many hip joint prostheses where the contact stresses can exceed 30 MPa. The major crux of the sustainability matter is that these materials are not natural to the human body and as simple as that is they cannot be adequately sustained. In economic terms the cost borne by the agencies and hospitals providing this implants climbs with innovation as well as by the year, this is of course a militating trend and most certainly not sustainable. (6) Possible solutions that can be addressed. It is not possible to avoid failure; recent work has focused on predictive tools to enable more accurate prediction so as to avoid catastrophic failure in vivo. [9]. The poor tribological properties of titanium alloys as compared to cobalt chromium alloys articulating against UHMWPE acetabular cups has prompted the use of surface treatments such as plasma vapour deposition coating of TiN and TiC, thermal treatments (nitriding, surface hardening), and ion implantation (N+) [49][9]. New materials are constantly being developed. Researchers recognised that wear debris will always be generated when two surfaces are in sliding contact. [9]. The challenge is developing engineering solutions to take care of wear debris rather than eliminating wear debris, a seemingly impossible task. [9]. Increasing hardness and fatigue resistance may only be a partial solution. [9]. Future advances may take one or more of the following routes: 1. Interpenetrating network composites: nanolaminate layer of interpenetrating network composites such as those found in nature have unique fracture resistance. [9]. Examples are seashells, which have been shown to give improved fracture resistance with unique wear characteristics [52][9]. The microstructure is made of a nano brick type arrangement of a ceramic phase sandwiched by an ultra-thin polymeric protein layer. [9]. Presumably, the small brick like ceramic components (often biodegradable) allow easy removal/dissolution, a concept that needs to be mimicked in engineering a biomaterial that has wear debris that is eco-compatible. [9]. By using a laminate concept, fracture toughness values as high as 16 MPa m can be achieved, for instance for boron carbide/aluminium laminates (Fig. 3). [9]. These laminates also have high flexural strength. Interpenetrating network composites such as those by bi-axial stretching of one crystalline phase (UHMWPE) and infiltrating with elastomeric polyurethane (PU) [53][9] to produce microlaminates has shown significant improvement in the strength and fracture toughness of an elastomeric composite membrane (less than 40 m) for biomedical applications (Fig. 4). [9]. Fig. 3. Fracture toughness versus specific flexural strength of some bioceramics and nanolaminates of metal matrix ceramic composites. [9]. Note the effect of laminates in improving both fracture toughness and flexural strength (after Saikaya and Aksay [52][9], courtesy of Springer-Verlag). Fig. 4. (a) Stress–strain behaviour of interpenetrating network composites of porous bi-axial UHMWPE and infiltrating with elastomeric PU to produce microlaminates with significantly improved mechanical properties [9]; (b) shows the cross-section view of the internal microstructures (after Teoh et al. [53])[9]. 2. A triplex phase composite consisting of a ductile phase, a hard phase and a lubricating phase that protects both articulating surfaces (e.g. Ti–TiC–graphite [54][9], which was designed with pure titanium providing the ductile damage resistant phase, titanium carbide the hard wearing phase, and the graphite the lubricating phase (Fig. 5)). [9]. 8 3. Engineered biomaterial surfaces and tribosystems that are able to trap/isolate wear debris and promote easy removal of such wear debris. [9]. Fig. 5. Microstructure of a Ti–TiC–graphite composite with improved wear resistant characteristics (after Teoh et al. [54])[9]. 9 References: (1) P.A. Dearnley; A brief review of test methodologies for surface-engineered biomedical implant alloys; Surface and Coatings Technology - Article in Press; Available online 30 December 2004. (2) http://www.bccresearch.com/archive/P211.html. (Assessed 28/04/05, 10:20 pm). (3) S. Ramakrishna, J. Mayer, E. Wintermantel and Kam W. Leong; Biomedical applications of polymer-composite materials: a review; Composites Science and Technology; Volume 61, Issue 9 , July 2001, Pages 1189-1224. (4) http://www.packereng.com/bioengineering.cfm. (Assessed 29/04/05; 9.03). (5) Clark M. Stanford and Galen B. Schneider; Functional behaviour of bone around dental implants; Gerodontology; Volume 21 Issue 2 Page 71 - June 2004. (6) http://www.mscsoftware.com/assets/EMDM_November.pdf. (Assessed 30/04/05; 7:19). (7) P. A. Dearnley, K. L. Dahm and H. Çimeno lu; The corrosion–wear behaviour of thermally oxidised CP-Ti and Ti–6Al–4V; Wear Volume 256, Issue 5 , March 2004, Pages 469-479. (8) http://www.i-sis.org.uk/EMIPAP.php. (Assessed 01/05/2005; 10.57). (9) S. H. Teoh; Fatigue of biomaterials a review; International Journal of Fatigue Volume 22, Issue 10 , November 2000, Pages 825-837.
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Comments Similarly for a nonuniformly distributed stress * = * (V): pf = ? (* - *u/*0)M dV/V0 V [1] Where *u is a threshold stress below which risk of rupture is zero. I just want to refer to the line i just pasted above: I normally get confused about rupture stress, with this my doubts are cleared. Thanks. It also happened that your surname is my username.
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Date of Posting (31/03/2009) 01/04/2009 00:00:00
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