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is the newsletter of PedicleScrew'd
Editor: Katt Morris
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Pedicle Screws that dissolve?

For over 10 years, companies abroad and in the U.S. have been working with a polymer substance that dissolves after a set time-frame has passed.

Bionx Implants develops, manufactures, and markets resorbable polymer implants for use in orthopedic surgery, urological stents, dentistry, and maxillo-facial surgery. These devices are made of substances that dissolve, making it unnecessary to do a second device-removing surgery.
For example, many times fractures require pins and screws for stabilization to allow proper healing. These pins and screws must later be removed, requiring a second surgery. Bionx Implants makes screws that slowly dissolve, obviating the need for removal.

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Whilst the use of metals can be traced for thousands of years, there wide spread application as surgical implants is limited to the past century. During the period of 1860 to 1900 whilst metalology was transforming rapidly from an empirical art into a highly complex technology orthopaedics was still in its infancy.
In this review of the role of metals in orthopaedic surgery, we shall look at the metals and their alloys that we currently use. We will also address current issues with regard to the systemic and remote side effects of the various biomaterials that are used, and the local response (such as osteolysis) to orthopaedic biomaterials.

The first metal alloy developed specifically for human use was "vanadium steel" in the early 1900's. In 1920 attempts to employ surgical implants were still hampered by the limitation of available materials. In 1924 Zierald published the study on the reaction of tissues to a variety of metals. Iron and steel, the most widely employed materials at the time, were noted to dissolve rapidly and to provoke erosion of adjacent bone. Substantial discolouration of tissues was observed around specimens of copper, nickel, embedded in bones. A problem existed, as the metals which did not produce discolouration, eg gold, silver, or pure aluminium, were all too soft or weak for most applications.
In 1926 18% chromium, 8% nickel stainless steel was introduced into surgical applications. This material was noted to be much more corrosion resistant in body fluids. This was stronger and more resistant to corrosion than the vanadium steel initially introduced by Sherman for his fracture fixation plates. Later in 1926, 18-8SMo stainless steel, which contained a small percentage of molybdenum, to improve the corrosion resistance in salt water, was introduced. This alloy became known as 316 stainless steel.
The next alloy to be introduced into orthopaedic practice was titanium and its alloys. In 1947 possible applications for titanium surgical implants were considered. The pure metal had shown excellent inertness in an environment of seawater, so that the corrosion resistance seemed likely to be good in the human environment. A few surgical implants were made and inserted into human subjects. Upon their removal, excellent corrosion resistance was confirmed. Subsequently, Maurice Down introduced a variety of titanium orthopaedic fracture devices such as plates and screws.
In the 1950's the carbon content of 316 stainless steel was reduced from 0.08% to 0.03% for better corrosion resistance and became known as 316L stainless steel.

The most important terms used to describe the characteristics of metals are;
1) Yield strength, or the stress necessary to produce plastic deformation in a material or permanently bend it;
2) The fatigue strength, the maximum load that a metal can withstand without fracturing when subjected to 10 million cyclic loads;
3) The ultimate tensile strength, the load that a metal can withstand in a single application without braking. When the ultimate tensile strength is exceeded, the metal will break.
The stress strain diagram shown in the figure below illustrates the effect of a single stress or load and clearly shows the progressive response to a larger load, but does not demonstrate the response to cyclic loading.
The alloys used in total joint components include: stainless steel, titanium, aluminium, vanadium, cast and forged cobalt chromium molybdenum, wrought cobalt chromium tungsten nickel and cobalt nickel chromium molybdenum. The metals can more simply be classified as the iron-based, the titanium-based, and the cobalt-based alloys. The characteristics of the metals have been standardised by the voluntary efforts of their producers, the device manufacturers, the material scientists, and the orthopaedic surgeons in the American Society for Testing of Metals (ASTM). There are no ASTM standards for the forging and casting process, the actual manufacturing process used to make components, or the performance. Also, there are no standard ASTM methods of testing for failure of a metal, and because there are a number of ways of testing for this characteristic, the fatigue data are difficult to interpret.
The manufacturing process can improve the strength characteristics of a metal considerably by minimising the defects (such as bubbles, voids, pieces of slag). Some inclusions are present in all metals and standards have been set for the maximum number acceptable. Too many large inclusions will weaken the component;
those on the surface will act as stress risers and provide areas for crevice corrosion. Galante, Laing and Lautenschlager have reported metallurgic defects of this type in some broken stems. Stem failures always cause suspicion of a defect in the design or in the metal, but the relatively low incidence of stem failures reported and the ability to demonstrate metallurgic defects of more than the maximum acceptable standard in most stems, suggest that the problem is actually technical not biomechanical. The ideal metal for a cemented component would have a high fatigue limit, yield strength, and ultimate tensile strength. Theoretically, a high modulus of elasticity (less elastic) may be considered advantageous because it would reduce the stress in the cement around the component and decrease the risk of cement failure but it would be disadvantageous because the bone may become so unloaded that disuse osteoporosis, or stress shielding, could develop, resulting in cement failure and producing loosening of the component.
Metals differ considerably in mechanical properties and chemical composition. Fracture of a total joint component starts in the region of highest tensile stress. A crack may be initiated and with additional loading, progress to complete fracture. The grain or crystal size of a metal is broadly indicative of its quality. In general, the larger the grain, the less the tensile strength of the metal; conversely, the smaller or finer the grain, the greater the toughness or strength. Heating metal to approximately its melting point decreases the grain size; forging decreases the grain size. Deforming a stainless steel stem increases the grain size primarily on the outer surface. HIPing (hot isostatic pressing) is a manufacturing procedure that simultaneously applies both heat and pressure to consolidate a part. It is used often to consolidate powder into a solid form and to reduce the porosity, thereby increasing the strength characteristic of a cast component.

Despite the great numbers of metals and alloys known to man, remarkably few warrant even preliminary consideration for uses implantable materials. The highly corrosive environment combined with the poor tolerance of the body to even minute concentrations of most metallic dissolution products eliminates from discussion all except certain materials based on iron, cobalt, nickel, titanium, tantalum, zirconium, silver, gold and the noble metals. Of this group tantalum and the noble metals do not have suitable mechanical properties for the construction of most orthopaedic tools and implants, while zirconium is in general too expensive.
All of the implantable metallic materials are used as alloys. An alloy is a metal made by taking a pure metal and adding other elements (the alloying elements) to it. A binary alloy has two components, a tertiary alloy has three, etc. The alloys are made from either wholly metallic components or contain some minor amounts of non-metallic elements such as carbon, sulphur, or oxygen. The precise composition of an alloy has profound effects on its phase structure and hence its properties.

The austentic stainless steels, especially Types 316 and 316L, are most widely used for implant fabrication. Stainless steel that has a low content of impurities and a passivated finish is entirely suitable for implantation in the human body. Forged stainless steel has a greater yield strength than cast stainless steels, but has a lower fatigue strength than other implant alloys. However, stainless steel is more ductile and more easily machined, and recent advancements have significantly enhanced its properties. Because a femoral component fracture with early designs, stainless steel is no longer used routinely, from the standpoint of erosion, biocompatability, and fatigue life, stainless steel is inferior to other super alloys. (However, it still may have applications in elderly patients in whom physical demands and life expectancy are limited, especially when cost is a major determinant.)
The only difference in composition between 316L and 316 stainless steel is the content of carbon. A wide range of properties exists depending on the heat treatment (annealing to obtain softer materials) or cold working (for greater strength and hardness). Even the 316L stainless steels may corrode inside the body under certain circumstances in a highly stressed and oxygen depleted region, such as contact under screws or fracture plates. Thus, stainless steels are suitable to use only in temporary implant devices, such as fractures plates, screws and hip nails.

There are basically two types of cobalt chromium alloys; one is the cobalt CoCrMo alloy, which is usually used to cast a product and the other is the CoNiCrMo alloy, which is usually wrought by (hot) forging. The castable CoCrMo alloy has been used for many decades in dentistry and recently, in making artificial joints. The wrought CoNiCrMo alloy is a relative newcomer now used for making the stems of prosthesis for heavily loaded joints such as the knee and hip.
Cobalt-based alloys are highly resistant to corrosion and especially to attack by chloride within crevice. As in all highly alloyed metals in the body environment, galvanic corrosion can occur, but to a lesser extent than in the iron-based alloys. Cobalt-based alloys are quite resistant to fatigue and to cracking caused by corrosion, and they are not brittle, since they have a minimum of 8% elongation. However, as is true of other alloys, cobalt based alloys may fail because of fatigue fracture (but less often than stainless steel stems).
The abrasive wear properties of the wrought CoNiCrMo alloy are similar to the cast CoCrMo alloy; however, the formula is not recommended for the bearing surface of joint prosthesis because of its poor frictional properties with itself or other materials. The superior fatigue and ultimate tensile strength of the wrought CoNiCrMo alloy make it suitable for the applications which require long service without fracture or stress fatigue. Such is the case for the stems of the hip joint prosthesis.
Both the cast and wrought alloys have excellent corrosion resistance.
The modulus of elasticity for the CrCo alloys does not change with the changes in their ultimate tensile strength. The values are higher than other materials such as stainless steels.
This may have some implications of different load transfer modes to the bone in artificial joint replacements, although the effect of the increased modulus on the fixation and longevity of the implants is not clear.

Attempts to use titanium for implant fabrication dates to the late 1930's. It was found that titanium was tolerated in cat femurs, as was stainless steel and vitalium (CoCrMo alloy). Titaniumís lightness and good mechano chemical properties are salient features for implant application. One titanium alloy (Ti6Al4V) is widely used to manufacture implants. The main alloying elements of the alloy are aluminium (5.5 - 6.5%) and vanadium (3.5 - 4.5%).
Whilst the strength of the titanium alloys varies from lower than to equal to that of 316 stainless steel, when compared by specific strength (strength per density), the titanium alloys excel any other implant material. Titanium nevertheless, has poor shear strength, making it less desirable for bone screws, plates and similar applications. Titanium also tends to gaul or seize when in sliding contact with itself or other metal.
Titanium-based alloys that have a high co-efficient of friction which can cause problems. Wear particles are formed in a piece of bone if a piece of bone rubs against the implant, or if two parts of an implant rub against one another. Therefore, implants of titanium upon titanium generally are not used as joint surfaces. Titanium derives its resistance to corrosion by the formation of a solid oxide layer. Under in vivo conditions, the oxide (TI02) is the only stable reaction product. The oxide layer forms a thin adherent film and passivates the material. Corrosion resistance mechanisms will be discussed later.

For orthopaedic applications of surgical implants fatigue failure is probably the primary concern. This occurs as a result of many applications of a stress lower than that required cause failure in a single application. Fatigue failure occurs at significantly lower life times as the applied stress increases. For some metals or alloys such as steels, a stress level is observed below which failure does not occur. This "fatigue limit" is empirically observed to be about 0.4 times the ultimate tensile strength.
In metals the crack usually initiates at the surface and grows slowly during the functional life of the specimen. The deformation tends to be localised within a few grains and develops into persistent slip bands which cannot be removed by polishing and etching. This crack may propagate in a zigzag fashion. Each stress cycle further growth occurs. Eventually when the residual section is sufficiently small and highly stressed it fails by conventional overload mechanisms. A number of factors effect fatigue life. Surface notches or holes severely effect fatigue strength. So that defects such as scratches or corrosion can reduce the strength by up to 20%.
Fatigue failure of surgical alloys is gradually becoming the principle cause of implant failure. With the introduction of total joint replacements, the prolonged implantation of metals has greatly increased. Since approximately 3 million flexions of the hip or spine per year is typical for the sedentary individual, the serious fatigue considerations for a Total Hip Replacement or Harrington rod for spinal immobilisation may readily be appreciated. The highest values for fatigue limits are for wrought cobalt chrome alloy while the purest grade of commercially pure titanium show the lowest value. Fatigue failure is usually associated with poor design, workmanship or handling. Attempts have been made to eliminate stress concentration such as crevices, corners or other irregularities. In practice however, even if the implant arrives in the theatre in pristine condition, scratches will inevitably occur during surgical implantation. The best hope for improvement rests on the possible introduction of a variety of fibre reinforced composite materials that show remarkable resistance to fatigue.

A major concern with all joint replacements is the degradation of the metals used (Black 1988). Wear and release of soluble products from implant materials results from the degradation of materials. In the case of alloys, soluble ions and compounds are released due to corrosion. Corrosion is defined as the unwanted chemical reaction of a metal with its environment, resulting in its continued degradation to oxides, hydroxides or other compounds.
To a certain extent metal ion release from alloys in the body always occurs, although the ions released are relatively negligible and generally this low level of contamination is not classed as corrosion. Corrosion occurs as a consequence of galvanic action or the breakdown of a passive protection film. An example of this is commonly found in the crevice that occurs between the screws and plate of a stainless steel fracture fixation device where the low O2 levels, reduction in PH, the build up of chloride ions and the formation of an anodic region results in corrosive attack of the alloy under the screw head.
Fretting of the morse tapers used for the attachment of the modular heads of THRís is a consequence of both wear and crevice corrosion and leads to the generation of a fine particulate debris (Matheson et al 1991).
These problems are well illustrated by looking at the problems faced by the Bone Tumour service at the Royal National Hospital. As a result of the extent of the large amounts of bone being excised during tumour surgery, many if not most patients require massive prosthesis to be implanted. Of particular concern with massive prosthesis is the wear and release of metallic ions from titanium alloy. The reason for this is that the shaft of the prosthesis and the intra medullary stem are made from titanium alloy, which can be worn either by rotation of the stem in the cement mantle or by soft tissue rubbing the part that replaces the bone (Lilley 1992). This leads to the generation of fine wear debris. This debris, which in part is derived from the titanium oxide layer which develops on the alloy surface or is composed of small titanium alloy particles, discolours the tissues such that on revision of massive prosthesis the interface around the shaft is often discoloured (Blunn 1991).
Patients with proximal tibial replacements sometimes present with the discolouration over their shin region which is caused by titanium wear debris. To reduce this wear the shafts of these replacements have over the last three years been coated with a layer of titanium nitride.
Systemic distribution of metal and UHMWPE particles from standard joint replacements to remote sites such as lymph nodes and spleen have been previously demonstrated. In a study by Albores - Saavedra et al (1994) six subjects with THRís undergoing prostatectomies and removal of lymph nodes for adenocarcinoma of the prostate or transitional cell carcinoma of the bladder, demonstrated discoloured lymph nodes with florid sinus histiocytosis characterised by large polygonal histiocytes (containing cobalt chromium or titanium particles) filling and expanding sinuses and interfolicular regions. A study by Case et al 1994 of 13 post mortem subjects with stainless steel and cobalt chromium implants, demonstrated dissemination of metal debris to local and distal lymph nodes, bone marrow, spleen and liver. In this study necrosis of lymph nodes was evident where there was most wear. A recent study has demonstrated the transfer of polyethylene wear particles from the implant site via the re lymphatic system to the lymph nodes of the contra lateral unoperated limb (Morawski 1995). More locally wear particles usually derived from the articular surface are able to invade the bone implant interface where they are phagocytosed by macrophages, which induces the release of various cytokines and this in turn stimulates the osteoclast. Distribution of wear debris and corrosion products from massive prosthesis can therefore be expected.
The corrosion products from metallic implants also are systemically distributed. Coleman et al 1973 demonstrated increased levels of chromium and cobalt in urine, serum, finger nails and hair in patients after metal on metal (cobalt, chromium, molybdenum) total hip replacements.
A more recent study has demonstrated an eight fold increase in chromium levels in serum from patients with long term metal on metal THRís compared with control groups (Jacobs 1996).
In a study of patients with cobalt chromium articulating against plastic components, elevated levels of chromium ions were detected in the serum of patients with modular implants and where the stem was cemented (Skipall 1996). This agrees with the finding of Karrholm (1994) and suggests that active dissolution mechanisms of fretting at either the stem-cement, head-neck interface or three body abrasion at the articulating surface are the dominant modes of metal ion release from THRís.
Some of the constituent metals used in prosthesis are needed by the body as trace elements. However, at higher concentrations these elements. Also have toxic effects. For example, in the 1970's it was discovered that dialysis patients from certain parts of the country accumulated aluminium derived from tap water in the cytoplasm of brain cells. The primary effect of this accumulation was damage to the brain which manifested itself as speech disturbances, myochronic jerks, motor apraxia, memory disturbances, personality changes, dementia and seizure disorders (Dialysis encephalopathy). The secondary effect of high levels of aluminium in dialysis patients is the development of osteomalacia where aluminium displaces the calcium in the osteoid resulting in an increased calcium level in the blood stream leading to an inhibition of parathyroid hormone. Aluminium rich neurofibrillar tangles have been reported in the cerebral cortex of deceased individuals suffering from Alzheimer dementia of note. Titanium alloy is composed of 6% aluminium. The role of metals derived from implants in immunological and oncogenic process is poorly understood. Metal ions are haptens and are capable of binding to proteins and forming antigens there is around a 10% incidence of sensitivity to chromium, cobalt and nickel in normal populations. Contact hypersensitivity reactions to metals in hip prosthesis had been described in the past and have been associated with high concentration of cobalt chromium and nickel in the tissues around implants. Metal on metal designs such as the McKee Farrar THR where relatively large amounts of metal debris were generated were prone to cause metal sensitisation (Elves 1975, Jones 1975). For titanium alloy there have been few reports of the constituent metals causing hypersensitivity reactions. Peters et al (1994) reported on sensitivity to titanium in one patient with a titanium pacemaker. A study by Laylaw et al 1991 demonstrated positive skin tests (to ointment contained titanium) in two out of five patients with failed THRís. The interface around these prosthesis was characterised by tissues containing titanium alloy derived wear debris accompanied by T-lymphocytes in the absence of B-lymphocytes suggesting sensitisation to titanium.
The effect of corrosion products derived from metal implants on the incidence of carcinoma is even more poorly understood. Ions of cobalt, chromium and nickel have been shown to be carcinogenic in animals.
The incidence of tumours around implant used for fracture fixation in animals is relatively common (Stevenson 1982). In man the relationship between implants and the development of carcinoma is very unclear. Malignant tumours at the implant site are extremely rare bearing in mind the number of individuals undergoing total joint replacements. Effective implants and corrosion products on the development of carcinomas at distant sites is even more uncertain. Concern has been expressed about the possible association between THR and malignancies of the lymphatic and haematopoietic systems. A recent study by Mathiesen et al 1995 of 1.6 million subjects with a cohort of 10, 785 individuals with THRís demonstrated no increased incidence of leukemia and lymphoma after total hip replacement.

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