Properties of the New Materials Used in Human Prosthetics and Orthotics

Transcription

1 Properties of the New Materials Used in Human Prosthetics and Orthotics Barbu D. Daniela Mariana; Barbu Gh. Ion Abstract During the last century, manufactured materials and devices have been developed to the point at which they can be used successfully to replace parts of living systems in the human body. These special materials, able to function in intimate contact with living tissue, with minimal adverse reaction or rejection by the body, are called biomaterials. Devices engineered from biomaterials and designed to perform specific functions in the body are generally referred to as biomedical devices or implants. This paper presents some general aspects of biomaterials used in orthosis and prosthesis manufacturing. Keywords: Biomaterials, Prosthesis, Orthosis, Eye Glasses. 1. INTRODUCTION Engineering materials, in many forms, have been implanted into the body since the introduction of aseptic surgery by Lister towards the end of the eighteenth century. Implant surgery has developed to the extent that implants are now used in most ranches of surgery and are becoming increasingly more sophisticated. Orthopedic surgery probably uses more implants than any other branch of medicine, fracture fixation being one of the oldest and most common applications. In recent years various joint replacements have been developed and, in addition to implants used in the musculo-skeletal system, a whole host of devices are used throughout the body for reconstructive, replacement, cosmetic and psychological reasons. The earliest successful implants were bone plates, introduced in the early 1900s to stabilize bone fractures and accelerate their healing. Advances in materials engineering and surgical techniques led to blood vessel replacement experiments in the 1950s, and artificial heart valves and hip joints were under development in the 1960s. As early as the first bone plate implants, surgeons identified material and design problems that resulted in premature loss of implant function, as evidenced by mechanical failure, corrosion and poor biocompatibility. Many implants have failed in the past and even now there are occasional failures. 1 Many of the failures are directly attributable to poorly chosen or badly used materials. For use in the body, a material must be completely non-toxic and any wear debris must also be non-toxic. In addition the material must behave satisfactorily with regard to all the properties listed below, and in some circumstances comply with further requirements: 1. adequate mechanical strength for both static and fatigue loading, in tension, compression and shear; 2. sufficient stiffness; 3. reasonable ductility; D.M. Barbu is with the Transilvania University of Brasov, Romania, phone: ; dbarbu@unitbv.ro. I. Barbu is with the Transilvania University of Brasov, Romania, phone: ; ibarbu@unitbv.ro. 4. sufficient hardness; 5. resistance to corrosion in the presence of body fluids, singly or in combination with other materials; 6. long term stability, both dimensional and chemical; 7. wear resistance in relation to mating parts; 8. capability of being sterilized. 2. METALS 2.1. Mechanical Strength It is obvious that an implant must have sufficient strength to withstand the various stresses encountered in service, making due allowance for any impact loading or fatigue. However, if the implant is much stronger than necessary (if the maximum stresses are low) then it will be heavier and larger than it needs to be. This can make the implantation more difficult for the surgeon and more traumatic for the patient and, in certain cases, even make the prosthesis less successful in service. So, the implant material should be used as efficiently as possible within the limits set by knowledge of the conditions of service of the implant Stiffness The stiffness requirement varies considerably depending upon the nature of the implant. In certain cases high stiffness or rigidity is required, as in a bone plate, whose purpose is to maintain the correct relationship of various sections of bone during the healing process. Movement of the bone at the fracture site has to be resisted and therefore materials with a high Young's modulus, such as stainless steel or chrome-cobalt alloy, should be used. Both these materials have modulus values of approximately 200 GPa. Titanium is another possibility in this situation but has a lower modulus of 110 GPa. Certain other situations require materials to have great flexibility and hence low values for the modulus of elasticity. Typical examples of this type are the integral finger hinge prosthesis and soft tissue replacement prostheses. This requirement of a low modulus of elasticity can be fulfilled by the use of a polymeric material. The polypropylene used in one type of finger prosthesis has a modulus of about 1 GPa. 423

2 2.3. Ductility and Brittleness Ductility is important in most loaded implants for the following reasons. Any prosthetic device will have minor flaws in its material structure, especially at the surfaces. In a stressed situation these flaws can lead to material failure, but this is much less likely in a ductile rather than a brittle material. The ability of a ductile material to deform plastically enables local stress concentrations to be relieved by yielding, thereby avoiding the propagation of the cracks. A brittle material is unable to yield locally and hence the load acting on the material is applied over a small area and extremely high local stresses ensue, leading to crack propagation and failure. In general, ductile materials are tougher and more able to withstand impact loads, because toughness is a measure of the energy needed to fracture a piece of material and this is a function of strength and elongation to break Hardness Hardness is perhaps the least important property of an implant material. Nevertheless, some consideration must be given to the degree of hardness required for a given prosthetic device. It is obvious that for a soft tissue implant a very soft material is required. In an articulating joint replacement, however, the wear rate is affected by the hardness of the mating materials and therefore, in this case, the hardness values must be known and controlled Corrosion Resistance The corrosion problem is associated almost exclusively with the metallic components of prosthesis. Most of the polymeric materials are affected relatively little by the body fluids, at least in the short term. The long term effects are discussed later in more important. However, all metals corrode in a hostile environment and the reaction is always electrochemical in nature. In some metals the early stages of corrosion build up an oxide layer at the surface, which eventually inhibits any further corrosion. The body offers a particularly hostile environment to metals and only those metals, which have the ability to form a protective oxide layer can be used in prostheses. Three such metals are titanium, chrome-cobalt alloy and, to a lesser extent, stainless steel (type 316). Metals can be arranged in series, which indicate the likelihood of corrosion caking place in service. The electrochemical series lists the normal electrode potentials of metals, usually in relation to hydrogen. While this series gives a general guide to the reactivity of various metals, it does not take into account the oxide film forming capability of these metals in any given electrolyte and it is more useful from the engineer's point of view to refer to the galvanic series in which metals are ranked in order of their relative reactivity in saline solutions. The following is a typical galvanic series, with the metals listed in order of increasing corrosion resistance: magnesium, zinc, low alloy steel, stainless steel 410 (active), stainless steel 316 (unpassivated), copper, nickel, silver, stainless 424 steel 410 (passive), stainless steel 316 (passive), titanium. It can be seen that the corrosion resistance of stainless steel is much improved by a process known as passivity. This involves the prosthesis being immersed in an oxidizing agent such as nitric acid, which thickens the oxide layer so that breakdown of this layer becomes more difficult. Galvanic corrosion occurs when there is a potential difference between the implant and the electrolyte. When an equilibrium is reached, i.e., when a sufficiently thick oxide film has developed, no further metal loss will take place. However, if the equilibrium is upset, corrosion will continue. Except in very rare and exceptional conditions this will happen when two dissimilar metals are in contact in the same electrolyte. Each metal will develop its own potential relative to the electrolyte and hence a potential difference will exist which will cause electrons to flow in the circuit. This will remove metal from that part of the implant, which has a relatively negative potential (i.e., the anode). Even slight differences in material composition within one component can lead to galvanic corrosion. Similarly, variations in solution concentration of the electrolyte over the different areas of the implant can lead to galvanic corrosion. The most important variation in the electrolyte is differential aeration, which induces corrosion in the metal due to the differing amounts of oxygen dissolved in the solution. Oxygen deficient regions become anodic and are attacked. In a multicomponent implant, the junctions between the various members can become deficient in oxygen and hence corrosion is more likely in these areas. Pitting of the metal is an example of attack due to oxygen deficiency. Corrosion is accelerated when the material is subjected to a mechanical stress and a hostile environment simultaneously. This is called stress corrosion cracking. This type of corrosion is always due to tensile stresses, either residual or applied, and some protection can be gained either by special surface coatings or by inducing residual compressive stresses into the material. Corrosion fatigue can be caused by any type of stress and the fatigue life of a material in a corrosive environment is always lower than it is in a noncorrosive environment. Although corrosion is a problem within the body, it is rare for corrosion of the currently acceptable biometals (austenitic stainless steel, chrome-cobalt alloy and titanium) to proceed to the extent where mechanical failure is a possibility. In implant surgery we have the unique circumstances in which a surgeon is utilizing materials under exceptionally stringent conditions and yet normally he has no idea of the properties of these materials. Severe cases of corrosion result either from the incorrect use of the materials or from the supply of material out of specification or in the wrong metallurgical condition. Galvanic corrosion occurs because surgeons sometimes use dissimilar metals in contact and crevice corrosion occurs because of a lack of awareness of the results of poor assembly. 3. PLASTICS

3 3.1. Long-term stability McHattie (1971) has discussed this aspect in detail. The problem can be stated simply - the stability or lack of it, of the material it self or the reaction to it induced in the body. In many ways the bodily environment is remarkably mild and commonplace. The fluid in which tissues are bathed resembles dilute sea water in analysis. Oxygen is freely available in normal tissues. However, by virtue of the same processes which permit such a wealth of chemical reaction to take place under these conditions in normal maintenance of life, the body becomes a remarkably hostile environment to foreign materials. The precise mechanisms by which degradation of implanted polymers take place are not yet known in detail but the extreme activity of hydrolyric and enzymatic reactions in the body suggest that only the most inert materials, free from chemical groups which can be hydrolysed, and resistant to oxidative attack, will be unaffected over long periods. And, unlike most living tissue, plastic implants have no self regenerating or repairing power -any changes or degradation of the implant provides an opportunity for further interaction with the host. There is a general tendency of plastics to initiate the formation of blood clots. While this propensity is made use of in certain situations, e.g., the deposition of fibrin in the interstices of arterial grafts, it is basically a serious disadvantage and can lead to the blockage or malfunction of devices such as heart valves or, with equally serious effects, fragments may break away and be carried through the vascular system and cause arterial block. Synthetic arteries of less than about one quarter of an inch in diameter cannot be used since they would be blocked by the initial clotting. Elimination of surface irregularities to provide the maximum smoothness, allied to design to provide streamline flow, seem to provide the best approaches at present. In the production of plastics, many additives are used to produce or overcome particular effects. The concentration of these plasticizers, stabilizers, antioxidants etc., may vary from a few parts per million up to a substantial proportion of the total weight of the finished product. In the past the presence of these additives has often been ignored and that, in combination with the failure to appreciate that the use of generic class names was quite inadequate for the proper characterization of a plastic, must account for many of the problems in obtaining reproducible experimental data. More importantly, the potentially toxic nature of many of these additives needs emphasis, especially as they may leach out of the implanted material. Such plastic materials cannot be regarded as safe unless all the ingredients have been identified and tested for toxicity. Furthermore, the precise quantity of each additive must be known Wear resistance The relative importance of each component will change as wear continues and modifications take place to the rubbing surfaces. Generally, abrasive wear will be most important initially, and probably most harmful overall, 425 but adhesive and fatigue components play an increasingly important role. The environmental conditions, the type of loading, the counter face roughness, and the manner in which the two surfaces move against each other all affect the wear process. Because of this, it will never be possible to simulate outside the body the precise working conditions of a joint in vivo and it is possible to make only estimates of the life expectancies of prostheses Sterilization One problem, which can present great practical difficulties is that of sterilization. The standard methods are steam at temperatures ranging from 120 to 140 C, ethylene oxide gas, gamma radiation, and certain chemical liquids. Some plastics will not stand steam temperatures and some absorb ethylene oxide (a chromic tissue irritant which can cause haemolysis), which is exceedingly difficult to remove. Sterilization by gamma irradiation of intensity 2.5 to 3 Mrad can cross linking and affect the mechanical properties of material. 4. APPLICATIONS 4.1. Prosthetics biomaterials Artificial joints consist of a plastic cup made of ultrahigh molecular weight polyethylene (UHMWPE), placed in the joint socket, and a metal (titanium or cobalt chromium alloy) or ceramic (aluminum oxide or zirconium oxide) ball affixed to a metal stem. Figure 1. Artificial knee joints are implanted in patients with a diseased joint to alleviate pain and restore function. After about 10 years of use, these artificial joints often need to be replaced because of wear and fatigueinduced delamination of the polymeric component. Institute engineers are developing improved materials to extend the lifetime of orthopaedic implants such as knees and hips. This type of artificial joint is used to replace hip, knee, shoulder, wrist, finger, or toe joints to restore function that has been impaired because of arthritis or other degenerative joint diseases or trauma from sports injuries or other accidents. Joint replacement surgery is performed on many people in the world. In most cases, it brings welcome relief and mobility after years of pain. Materials and design engineers must consider the physiologic loads to be placed on the implants, so they can design for sufficient structural integrity. Material choices also must take into account biocompatibility

4 with surrounding tissues, the environment and corrosion issues, friction and wear of the articulating surfaces, and implant fixation either through osseointegration (the degree to which bone will grow next to or integrate into the implant) or bone cement. In fact, the orthopedic implant community agrees that one of the major problems plaguing these devices is purely materials-related: wear of the polymer cup in total joint replacements. Although the wear problem is one of materials, it plays out as a biological disaster in the body. Any use of the joint, such as walking in the case of knees or hips, results in cyclic articulation of the polymer cup against the metal or ceramic ball. Due to significant localized contact stresses at the ball/socket interface, small regions of prosthesis tend to adhere to the metal or ceramic ball. During the reciprocating motion of normal joint use, fibrils will be drawn from the adherent regions on the polymer surface and break off to form nanometer-sized wear debris. This adhesive wear mechanism, coupled with fatigue-related delaminating of the prosthesis (most prevalent in knee joints), results in billions of tiny polymer particles being shed into the surrounding sensorial fluid and tissues. The biological interaction with small particles in the body then becomes critical. The body s immune system attempts, unsuccessfully, to digest the wear particles (as it would a bacterium or virus). Enzymes are released that eventually result in the death of adjacent bone cells, or osteolysis. Over time, sufficient bone is resorbed around the implant to cause mechanical loosening, which necessitates a costly, and painful implant replacement, or revision. Since the loosening is not caused by an associated infection, it is termed aseptic. The average life of a total joint replacement is 8-12 years - even less in more active or younger patients. Because it is necessary to remove some bone surrounding the implant, generally only one revision surgery is possible, thus limiting current orthopedic implant technology to older, less active individuals. A relatively recent incident in the biomedical device field serves to illustrate the importance of materials choice and engineering on implant performance. The temporomandibular joint (TMJ) provides all jaw mobility and is crucial for chewing, talking, and swallowing. This joint can deteriorate from disease or trauma, which, in severe cases, necessitates replacement by an artificial joint. For many years, less than optimum technologies existed for TMJ implants. In the late 1970s, a TMJ replacement using polytetraflouro-ethylene (PTFE) as the bearing counterface was invented, and, in 1983, the inventors received FDA approval to market the PTFE implant, which was called the Interpositional Implant (IPI). In theory, PTFE would seem an appropriate choice for an implant material, as it exhibits a low coefficient of friction and has been used extensively as a bearing surface in other engineering applications. Materials engineers know that the reason PTFE exhibits such a low coefficient of friction is that a thin film of the material is continuously transferred onto the 426 opposing bearing surface. Although this transfer film acts as a lubricant, it also, by virtue of its formation, subjects the material to an adhesive wear mechanism. In the case of the PTFE TMJ implants, surrounding tissues quickly became overwhelmed by wear debris, and the immune system response resulted in osteolysis, causing massive destruction of the joint and surrounding tissues. For those people who received the implants, this was truly a tragedy; many suffered severe facial deformities, and most experienced unbearable pain and were no longer able to chew, swallow, or sleep. Figure 2. The wrist joint is a complicated one, allowing flexion, extension, adduction, and abduction, primarily through the radiocarpal joint. Though not subject to the severe wear problems encountered in other orthopedic implants, artificial wrist joints will benefit from improved biomaterials, particularly those designed to incorporate biological factors to enhance bone attachment, as well as materials with improved mechanical integrity and corrosion resistance. Work at designer is addressing the wear problem in UHMWPE total joint prostheses. They are studying the wear process and biological responses to wear debris. Results of these studies have led to novel ideas for materials modification and development. It is also developing new composite materials to defeat the fatigue-induced delaminating observed in the UHMWPE component of knee implants. Studies of wear debris extracted from actual tissue samples of patients whose implants failed as a result of aseptic loosening have generated significant information regarding wear particle size, shape, and surface morphology. Institute scientists were the first to use the atomic force microscope (AFM) to produce detailed, high-resolution images of wear particles Active biomaterials When a man-made material is placed in the human body, tissue reacts to the implant in a variety of ways depending on the material type. Therefore, the mechanism of tissue attachment depends on the tissue response to the implant surface. In general, materials can be placed into three classes that represent the tissue response they elicit: inert, bioresorbable, and bioactive. Inert materials such as titanium, UHMWPE, and alumina (Al 2 O 3 ) are nearly chemically inert in the body and exhibit minimal chemical interaction with adjacent tissue. A fibrous tissue capsule will normally form around inert implants. Tissue attachment with inert materials can be through tissue growth into surface irregularities, by bone cement, or by press fitting into a defect. This morphological fixation is not ideal for the

5 long-term stability of permanent implants and often becomes a problem with orthopaedic and dental implant applications. Bioresorbable materials, such as tricalcium phosphate and polylactic-polyglycolic acid copolymers, are designed to be slowly replaced by tissue (such as bone) or for use in drug-delivery applications. Certain glasses, ceramics, and glass-ceramics that contain oxides of silicon, sodium, calcium, and phosphorus (SiO 2, Na 2 O, CaO, and P 2 O 5 ) have been shown to be the only materials known to form a chemical bond with bone, resulting in a strong mechanical implant/bone bond. These materials are referred to as bioactive because they bond to bone (and in some cases to soft tissue) through a time-dependent, kinetic modification of the surface triggered by their implantation within living bone. In particular, an ionexchange reaction between the bioactive implant and surrounding body fluids results in the formation of a biologically active hydrocarbonate apatite (calcium phosphate) layer on the implant that is chemically and crystallographically equivalent to the mineral phase in bone. This equivalence is responsible for the relatively strong interfacial bonding. Although bioactive materials would appear to be the answer to biomedical implant fixation problems, available bioactive glasses are not suitable for loadbearing applications, and so are not used in orthopaedic implants. In fact, their use for other implants, even some dental applications, is limited because they have a low resistance to crack growth. However, there are stronger ceramic materials, crystalline in structure that is not as bioactive. Recent work has used ion beam surface modification to change the atomic structure and chemistry at the surface of these crystalline ceramics to allow the material to react (ion exchange) upon implantation Eyeglasses lenses materials BK7 is one of the most common borosilicate crown glasses used for visible and near infrared optics. Its high homogeneity, low bubble content, and straightforward manufacturability make it a good choice for transmissive optics. The transmission range for BK7 is nm. It is not recommended for temperature sensitive applications, such as precision mirrors. to 170 nm, and nonbirefringent properties make it ideal for deep UV transmissive optics. Material for IR use has grown using naturally mined fluorite, at much lower cost. CaF 2 is sensitive to thermal shock, so care must be taken during handling. Figure 4. Optical transmission of CaF 2 optical material Crystal Quartz is a positive uniaxial birefringent single crystal grown using a hydrothermal process. It has good transmission from the vacuum UV to the near infrared. Due to its birefringent nature, crystal quartz is commonly used for wave plates. Figure 5. Optical transmission of Crystal Quartz material UV grade fused silica is synthetic amorphous silicon dioxide of extremely high purity. This noncrystalline, colorless silica glass combines a very low thermal expansion coefficient with good optical qualities, and excellent transmittance in the ultraviolet. Transmission and homogeneity exceed those of crystalline quartz without the problems of orientation and temperature instability inherent in the crystalline form. Fused silica is used for both transmissive and reflective optics, especially where high laser damage threshold is required. Figure 3. Optical transmission of BK7 optical material Calcium Fluoride (CaF 2 ) is a cubic single crystal material grown using the vacuum Stockbarger technique with good vacuum UV to infrared transmission. CaF 2 s excellent UV transmission, down 427 Figure 6. Optical transmission of UV Fused Silica material Magnesium Fluoride (MgF 2 ) is a positive birefringent crystal grown using the vacuum Stockbarger technique with good vacuum UV to infrared transmission. It is typically oriented with the c axis parallel to the optical axis to reduce birefringent effects. High vacuum UV transmission, down to 150 nm, and its proven use in fluorine environments make it ideal for lenses, windows, and polarizers for Excimer lasers. MgF 2 is resistant to thermal and mechanical shock.

6 Abbe No. V d Coef. of Therm.Exp. (10-6 / C) Table 1. Optical properties of materials Conductiv. (W/m C) Heat Capac. (J/gm C) Density at 25 C (gm/cm 3 ) Knoop Hardn. (kg/mm 2 ) Young s Mod. (GPa) BK SF UV Fused Silica CaF MgF c axis 21 c axis 8.48 c axis 30 c axis Crystal Quartz c axis 10.4 c axis 97 c axis c axis 6.2 c axis 76.5 c axis Pyrex Zerodur ± less deposit resistant. Figure 7. Optical transmission of MgF 2 optical material Pyrex is a borosilicate glass with a low coefficient of thermal expansion. It is mainly used for nontransmissive optics, such as mirrors, due to its low homogeneity and high bubble content. Zerodur is a glass ceramic material that has a coefficient of thermal expansion approaching zero, as well as excellent homogeneity of this coefficient throughout the entire piece. This makes Zerodur ideal for mirror substrates where extreme thermal stability is required. Zerodur should not be used for transmissive optics due to inclusions in the material Contact lenses materials Fitting soft and rigid gas permeable lenses is a complex topic. Fitting concepts vary greatly and contact lens specialists need to consider the various measurement technology required as well as material selection, wear schedule and replacement modality for each individual patient. The following is an overview of what contact lens fitters need to know about both soft and oxygen permeable (GP), the industry s newly adopted term for rigid gas permeable, contact lenses before fitting them. A Dk measurement quantifies the oxygen permeability of all contact lens materials, relative to a diffusion coefficient and how much gas can pass through the material, known as solubility. The oxygen permeability of a material can be affected by environmental factors such as oxygen concentration, temperature and barometric pressure. A Dk/L measurement represents a calculated value of oxygen transmissibility, where L refers to the thickness of a given lens prescription, so that thick minus lens edges and thick plus lens centres decrease the transmissibility proportionately. It s important to note that numerous methods are employed to measure Dk s. The FDA categorizes soft lens materials into four groupings based on their water content (above or below 50 percent) and electrostatic properties. In general, as soft and GP increase in oxygen permeability, they become CONCLUSION To conclude this paper, the particular advantages and limitations of ultra-high molecular weight polyethylene, the most used material for prostheses, are: 1. It is the simplest polymer with no oxidizable side groups. 2. It has very long polymer chains, which give it excellent mechanical properties and good wear resistance. 3. It has a low coefficient of friction against steel. 4. It does not appear to deteriorate chemically or mechanically, or swell appreciably in the body. 5. It requires few additives. 6. It is nontoxic and the wear debris is nontoxic. 7. It can be sterilized by gamma irradiation. High molecular weight polyethylene has been used successfully for many years and that itself is a very significant factor in its favour. However, there are indications of three areas where problems might arise: the possibility of long-term environmental deterioration, the inevitable wear and creep. Evidence has been presented (Atkinson, 1980) that creep is the principal problem, at least in the initial stages, and one method of reducing this has been discussed. After long periods in service, wear will undoubtedly occur and there is also the possibility that enzymes eventually lead to partial degradation of the polymer (Williams, 1979), but the indications are that both processes, if they do occur, occur very slowly. 6. REFERENCES [1] Barbu, D.M., Tehnologii de montaj şi adaptare ochelari, Editura Univ.Transilvania Braşov, 2003, ISBN [2] Dumbleton, J. (1981). Tribology of natural and artificial joints, Elsevier Scientific, London; [3] Ludema, K. (1991). Tribological modeling for mechanical designers, ASTM, Philadelphia; [4] Stolarski, T.A. (1990). Tribology in machine design, Heinemann Newnes, Oxford; [5] Weber, J.; Humprey, F.; Silver, A., Modern Ophthalmic Frame Materials, Marchon, 1997; [6] Hostetter, T., Soft Contacts: The Hard Facts, Jobson Publishing L.L.C., 2001; [7]