Biomaterial

biomaterial is used to make devices to replace a part or a function of the body in a safe, reliable, economic, and physiologically acceptable manner. a variety of devices and materials are used in the treatment of disease or injury. commonplace examples include sutures, needles, catheters, plates, tooth fillings, etc. a biomaterial is a material of synthetic as well as of natural origin in contact with tissue, blood, and biological fluids, and intended for use for prosthetic, diagnostic, therapeutic, and storage applications without adversely affecting the living organism and its components. according to these definitions, one must have a vast field of knowledge or collaborate with different specialties in order to develop and use biomaterials in medicine and dentistry. the uses of biomaterials include replacement of a body part, which has lost function due to disease or trauma, to assist in healing, to improve function, and to correct abnormalities.


uses of biomaterials:

problem area examples
replacement of diseased or damaged part -artificial hip joint, kidney dialysis machine
assist in healing - sutures, bone plates, and screws
improve function - cardiac pacemaker, intraocular lens
correct cosmetic problem -augmentation mammoplasty, chin augmentation
aid to diagnosis - probes and catheters
aid to treatment - catheters, drains

biomaterials are governed by the interaction between the material and the body specifically, the effect of the body environment on the material and the effect of the material on the body. it should be evident from any of these perspectives that most current applications of biomaterials involve structural functions, even in those organs and systems, which are not primarily structural in their nature, or very simple chemical or electrical functions. complex chemical functions such as those of the liver and complex electrical or electrochemical functions such as those of the brain and sense organs cannot be carried out by biomaterials at this time.




classification of biomaterials:

biomaterials can broadly be classified as: i) biological biomaterials and ii)
synthetic biomaterials. biological materials can be further classified into soft and hard tissue types. in the case of synthetic materials, it is further classified into: a) metallic b) polymeric c) ceramic and d) composite biomaterials.

1. biological materials:

- soft tissue: skin, tendon, pericardium, cornea

- hard tissue: bone, dentine, cuticle

2. synthetic biomedical materials:

- polymeric: ultra high molecular weight polyethylene(uhmwpe),
polymethylmethacarylate (pmma), polyethyletherketone (peek), silicone, polyurethane (pu), polytetrafluoroethylene (ptfe).

- metallic: stainless steel, cobalt-based alloy (co-cr-mo), titanium alloy (ti-al-v), gold, platinum.

- ceramic: alumina (a1 2o3), zirconia (zro2),carbon, hydroxylapatite [ca10(po4)6 (oh)2], tricalcium phosphate [ca3(po4)2], bioglass [na2o( cao)(p2o3)(sio2)],
calcium aluminate [ca(a12o4)].

- composite: carbon fiber (cf)/peek, cf/uhmwpe, cf/pmma, zircon /silica.

requirements of biomaterials:

biomaterials must have special properties that can be tailored to meet the
needs of a particular application - this is an important concept to bear in
mind. for example, a biomaterial must be biocompatible, non-carcinogenic, corrosion-resistant, and has low toxicity and wear. however, depending on the application, differing requirements may arise. sometimes these requirements can be completely opposite. in tissue engineering of the bone, for instance, the polymeric scaffold needs to be biodegradable so that as the cells generate their own extracellular matrices, the polymeric biomaterial will be completely replaced over time with the patient s own tissue. in the case of mechanical heart valves, on the other hand, we need materials that are biostable, wear-resistant, and which do not degrade with time.
generally, the requirements of biomaterials can be grouped into four broad categories:

1-biocompatibility:

biocompatibility involves the acceptance of an artificial implant by the surrounding tissues and by the body as a whole. biocompatible materials do not irritate the surrounding structures, do not provoke an abnormal inflammatory response, do not incite allergic or immunologic reactions, and do not cause cancer. other compatibility characteristics that may be important in the function of an implant device made of biomaterials include (1) adequate mechanical properties such as strength, stiffness, and fatigue properties (2) appropriate optical properties if the material is to be used in the eye, skin, or tooth and (3) appropriate density.
sterilizability, manufacturability, long-term storage, and appropriate engineering design are also to be considered.
the failure modes may differ in importance as time passes following the implant surgery. for example, consider the case of a total joint replacement in which infection is most likely soon after surgery, while loosening and implant fracture become progressively more important as time goes on. failure modes also depend on the type of implant and its location and function in the body. for example, an artificial blood vessel is more likely to cause problems by inducing a clot or becoming clogged with thrombus than by breaking or tearing mechanically.

2-sterilizability:

the material must be able to undergo sterilization. sterilization techniques include dry heat, gamma, gas (ethylene oxide (eto)) and steam autoclaving. some polymers such as polyacetal will depolymerize and give off the toxic gas formaldehyde when subjected under high-energy radiation by gamma. these polymers are thus best sterilized by eto.
sterilizability of biomedical polymers is an important aspect of the properties because polymers have lower thermal and chemical stability than other materials such as ceramics and metals consequently, they are also more difficult to sterilize using conventional techniques.
in dry heat sterilization, the temperature varies between 160 and 190°c. this is above the melting and softening temperatures of many linear polymers such as polyethylene and pmma. in the case of polyamide (nylon), oxidation will occur at the dry sterilization temperature although this is below its melting temperature. the only polymers, which can safely be dry, sterilized are ptfe and silicone rubber.
steam sterilization (autoclaving) is performed under high steam pressure at relatively low temperature (125–130°c). however, if the polymer is subjected to attack by water vapor, this method cannot be employed. pvc, polyacetals, pe (low-density variety), and polyamides belong to this category.
chemical agents such as ethylene and propylene oxide gases and phenolic and hypochloride solutions are widely used for sterilizing polymers since they can be used at low temperatures. chemical agents sometimes cause polymer deterioration even when sterilization takes place at room temperature. however, the time of exposure is relatively short (overnight), and most polymeric implants can be sterilized with this method.
radiation sterilization using the isotopic 60co can also deteriorate polymers since at high dosage the polymer chains can be dissociated or cross-linked according to the characteristics of the chemical structures, as shown in table 1 in the case of pe, at high dosage (above 106 gy) it becomes a brittle and hard material. this is due to a combination of random chain scission cross-linking. pp articles will often discolor during irradiation giving the product an undesirable color tint but the more severe problem is the embrittlement resulting in flange breakage, luer cracking, and tip breakage. the physical properties continue to deteriorate with time, following irradiation. these problems of coloration and changing physical properties are best resolved by avoiding the use of any additives that discolor at the sterilizing dose of radiation.


table 1 effect of gamma irradiation on polymers that could be cross-linked or degraded.

cross-linking polymers degradable polymers
polyethylene
polypropylene
polystyrene
polyarylates
polyacrylamide
polyvinylchloride
polyamides
polyesters
polyvinylpyrrolidone
polymethacrylamide
rubbers
polysiloxanes
polyvinylalcohol
polyacroleine polyisobutylene
poly-??methylstyrene
polymethylmetacrylate
polymethacrylamide
polyvinylidenechloride
cellulose and derivatives
polytetrafluoroethylene
polytrifluorochloroethylene


3-functionability:

the functionability of a medical device depends on the ability of the material to be shaped to suit a particular function. the material must therefore be able to be shaped economically using engineering fabrication processes. the success of the coronary artery stent - which has been considered the most widely used medical device — can be attributed to the efficient fabrication process of stainless steel from heat treatment to cold working to improve its durability.

4 -manufacturability:

it is often said that there are many candidate materials that are biocompatible. however it is often the last step, the manufacturability of the material, that hinders the actual production of the medical device.

performance of biomaterials:

the success of biomaterials in the body depends on factors such as the material properties, design, and biocompatibility of the material used, as well as other factors not under the control of the engineer, including the technique used by the surgeon, the health and condition of the patient, and the activities of the patient. if we can assign a numerical value f to the probability of failure of an implant, then the reliability can be expressed as:
r = 1 ? f
if, as is usually the case, there are multiple modes of failure, the total reliability rt is given by the product of the individual reliabilities r1 = (1 ? f1), etc.
rt = r1 • r2 • • • rn
consequently, even if one failure mode such as implant fracture is perfectly controlled so that the corresponding reliability is unity, other failure modes such as infection could severely limit the utility represented by the total reliability of the implant. one mode of failure which can occur in a biomaterial, but not in engineering materials used in other contexts, is an attack by the body’s immune system on the implant.
another such failure mode is an unwanted effect of the implant upon the body for example, toxicity, inducing allergic reactions, or causing cancer.




surface modifications for improving biocompatability:

prevention of thrombus formation is important in clinical applications where blood is in contact such as hemodialysis membranes and tubes, artificial heart and heart–lung machines, prosthetic valves, and artificial vascular grafts. in spite of the use of anticoagulants, considerable platelet deposition and thrombus formation take place on the artificial surfaces.
heparin, one of the complex carbohydrates known as mucopolysaccharides or glycosaminoglycan, is currently used to prevent formation of clots. in general, heparin is well tolerated and devoid of serious consequences. however, it allows platelet adhesion to foreign surfaces and may cause hemorrhagic complications such as subdural hematoma, retroperitoneal hematoma, gastrointestinal bleeding, hemorrhage into joints, ocular and retinal bleeding, and bleeding at surgical sites. these difficulties give rise to an interest in developing new methods of hemocompatible materials.
many different groups have studied immobilization of heparin on the polymeric surfaces, heparin analogs and heparin–prostaglandin or heparin–fibrinolytic enzyme conjugates. the major drawback of these surfaces is that they are not stable in the blood environment. it has not been firmly established that a slow leakage of heparin is needed for it to be effective as an immobilized antithrombogenic agent if not, its effectiveness could be hindered by being “coated over” with an adsorbed layer of more common proteins such as albumin and fibrinogen.
-albumin-coated surfaces have been studied because surfaces that resisted platelet adhesion in vitro were noted to adsorb albumin preferentially.
-fibronectin coatings have been used in in vitro endothelial cell seeding to prepare a surface similar to the natural blood vessel lumen.
-algin-coated surfaces have been studied due to their good biocompatibility and biodegradability.
- plasma gas discharge and corona treatment with reactive groups introduced on the polymeric surfaces have emerged as other ways to modify biomaterial surfaces.
-hydropinghobic coatings composed of silicon- and fluorine-containing polymeric materials as well as polyurethanes have been studied because of the relatively good clinical performances of silastic, teflon, and polyurethane polymers in cardiovascular implants and devices. polymeric fluorocarbon coatings deposited from a tetrafluoroethylene gas discharge have been found to greatly enhance resistance to both acute thrombotic occlusion and embolization in small diameter dacron grafts.
-hydropinghilic coatings have also been popular because of their low interfacial tension in biological environments.
-hydrogels as well as various combinations of hydropinghilic and hydropinghobic monomers have been studied on the premise that there will be an optimum polar-dispersion force ratio which could be matched on the surfaces of the most passivating proteins. the passive surface may induce less clot formation.
-polyethylene oxide coated surfaces have been found to resist protein adsorption and cell adhesion and have therefore been proposed as potential “blood compatible” coatings.
- saline perfusion method another way of making antithrombogenic surfaces, which is designed to prevent direct contacts between blood and the surface of biomaterials by means of perfusing saline solution through the porous wall which is in contact with blood. it has been demonstrated that the adhesion of the blood cells could be prevented by the saline perfusion through pe, alumina, sulfonated/nonsulfonated ps/sbr, eptfe (expanded polytetrafluoroethylene), and polysulfone porous tubes.

general physical and chemical methods to modify the surfaces of polymeric biomaterials are listed in table 2.

table 2 physical and chemical surface modification methods for polymeric biomaterials.
to modify blood compatibility silicon containing block copolymer additive
plasma fluoropolymer deposition
plasma siloxane polymer deposition
radiation-grafted hydrogels
chemically modified polystyrene for heparin-like activity
to influence cell adhesion and growth oxidized polystyrene surface
ammonia plasma-treated surface
plasma-deposited acetone or methanol film
plasma fluoropolymer deposition
to control protein adsorption surface with immobilized polyethyelenglycol
affinity chromatography particulates
surface cross-linked contact lens
to improve lubricity plasma treatment
radiation-grafted hydrogels
interpenetrating polymeric networks
to improve wear resistance and
corrosion resistance ion implantation
diamond deposition
anodization
to alter transport properties plasma deposition (methane, fluoropolymer, siloxane)

to modify electrical characteristics plasma deposition
solvent coatings


mechanical properties of biomaterials:
the mechanical properties of a biomaterial can best be described by its modulus of elasticity, ultimate tensile strength, elongation to failure, and fracture toughness.
• modulus of elasticity describes the stiffness of the material and is usually obtained from the slope of a stress-strain diagram.
• ultimate tensile strength describes the ability of the material to withstand a load before it fails.
• elongation to failure describes how much strain the material can bear before it fails.
• fracture toughness is an important measurement of the material s resistance to crack propagation.
figures 1 (a) to (d) show the comparisons amongst different classes of biomaterial with respect to the four properties mentioned above. it can be seen that metals are generally very stiff and have high fracture toughness. in sharp contrast to the metals are the polymers, which have low stiffness and fracture toughness. however the polymers have high elongation to failure. the high stiffness of metals, on the other hand, can be a disadvantage since this can give rise to "stress shielding" in bone fracture repair. stress shielding is a phenomenon where bone loss occurs when a stiffer material is placed over the bone. bone responds to stresses during the healing process. since the stress is practically shielded from the bone, the density of the bone underneath the stiffer material decreases as a result.