Corrosion in Body Fluids

2.27 Corrosion in Body Fluids D. J. Blackwood Department of Materials Science and Engineering, National University of Singapore, 9 Engineering Drive 1, Singapore 117576, Singapore

ß 2010 Elsevier B.V. All rights reserved.

2.27.1 2.27.2 2.27.3 2.27.4 2.27.5 2.27.5.1 2.27.5.2 2.27.5.3 2.27.5.4 2.27.5.5 2.27.5.6 2.27.5.7 2.27.5.8 2.27.5.9 2.27.6 2.27.6.1 2.27.6.2 2.27.6.3 2.27.6.4 2.27.6.5 2.27.6.6 2.27.6.7 2.27.7 References

Introduction Historical Development Health Effects Related to Corrosion in Body Fluids Environments Encountered in Biomedical Applications Metals and Alloys Used in Biomedical Applications Titanium and Titanium Alloys Cobalt–Chromium–Molybdenum Alloys Stainless Steels Nickel–Titanium Alloy Porous Materials and Metallic Foams Magnesium Alloys Rare Earth Magnets Dental Amalgams Titanium Nitride Coatings Corrosion Types Encountered in Biomedical Applications General Corrosion Pitting Corrosion Crevice Corrosion SCC and Hydrogen Embrittlement Corrosion Fatigue Fretting Corrosion and Wear Galvanic Corrosion Conclusions

Glossary Biocompatibility The extent to which an implanted material elicits an immune response in a host. Carcinogen Cancer causing. Cardiac Related to the function of the heart. Cytotoxicity Toxic towards cells. Extracellular The space outside the plasma membranes of cells and occupied by fluid; literally means outside the cell as opposed to intracellular which is inside the cell. Hematoma A collection of blood, usually partially clotted, that results from the breakage of veins or blood vessels. Inflammatory response Part of the human body’s initial response to injury or infection. In vitro Experiments conducted outside the body; literally means in-glass.

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1309 1309 1310 1311 1313 1313 1314 1314 1314 1315 1315 1316 1316 1316 1316 1316 1317 1317 1317 1318 1318 1319 1319 1320

In vivo Data collected from materials implanted within a live subject, human or animal. Orthodontic An area of dentistry concerned with treatment of the inability to bite correctly, due to the misalignment of teeth, irregular tooth growth or disproportionate relationship between the upper and lower jaws. Oral lesions Abnormal tissue in or around the mouth, usually caused by disease or trauma. Oral mucosa Mucous membranes of the mouth. Osteoporosis Thin or porous bones caused by lack of calcium or stress shielding by load-bearing implants. Prosthesis An artificial device that replaces a missing body part. Pathological changes The human body’s initial response to injury or infection. Pulmonary Related to the function of the lungs.

Corrosion in Body Fluids

Serum Blood plasma from which clotting factors have been removed. Shape-memory alloy A material that, after it has been deformed, regains its original geometry by heating, due to temperature-dependent martensitic phase transformation. Sherman plates Metal plates used to hold fractured bones in place to allow them to heal; typically the plate is screwed to the bone.

Abbreviations ASTM American Society for Testing and Materials FCC Face Centered Cubic ISO International Standards Organization L Low carbon PREN Pitting resistance equivalent number ¼%Cr þ (3.3%  Mo) þ (16  %N) SCC Stress corrosion cracking SHE Standard hydrogen electrode VM Vacuum melted

2.27.1 Introduction Corrosion is of concern to the biomedical industry for two reasons. The first is common to all industries, being ‘how will a device’s lifetime/performance be impacted by corrosion?’ However, the second concern is specific to the biomedical industry, being ‘will the metallic ions that leach out of the device build up to levels sufficient to harm the patient, for example, by causing tumors to develop?’ There is little known about what represents longterm safe limits for metallic ion concentrations in the body, and therefore, these could well be exceeded at corrosion rates that are insignificant with respect to the physical performance of the implant. Corrosion problems in dental applications are more common, mainly due to the high acidity and chloride contents of many foodstuffs. Fortunately, fixtures in the oral cavity are readily accessible for repair or replacement, but the toxicity of the metals leaching out is a major concern. In most industries, corrosion is typically controlled by either coatings or altering the local environment. Unfortunately, until recently, coatings were of limited use for protecting surgical implants, since many of these (especially orthopedic devices) are

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subjected to wearing and abrasion processes. Likewise, the environment within the human body is essentially fixed, and therefore, it is not possible to lower the temperature or raise the pH. As a result, the only available method of reducing corrosion rates within the human body has been to fabricate the biomedical devices from a corrosion-resistant material. Nevertheless, since the late 1970s, this approach has been extremely successful, at least with respect to extending the lifetime of biomedical devices, thanks to the development of a range of corrosion-resistant alloys that have reduced the number of failures to extremely low levels. Besides, many of the few corrosion failures that still occur can be traced either to poor quality control or to an unexpected and unusually aggressive local environment around the implant, due to pathological changes in the surrounding tissue as this reacts to the surgical procedure. The remaining early failures of surgical implants are due to either fatigue or fretting, which may or may not be accelerated by corrosion. Nevertheless, the lack of consensus over what represents safe levels for metallic ion concentrations within the body, or even which metals are toxic, means that the concerns about extended exposure to even very low levels of corrosion products resulting in medical complications remain. Furthermore, average life expectancies are increasing and the average age of patients receiving implants is decreasing. Ironically, this is both in part due to the modern popularity of physical sports, which place a large strain on joints, and therefore, the required performance lifetime of devices, and with it, the likelihood of both the corrosion-related failures and health problems instigated by elevated metallic levels is increasing. Additionally, a number of advanced materials, such as shape-memory alloys, are being introduced into the biomedical industry often with very little prior thought given to corrosion protection.

2.27.2 Historical Development The earliest use of metallic materials in dental applications was over 4000 years ago in ancient Egypt; although initially for aesthetic purposes, by about 700 BC, this had developed to the level of using gold bridgeworks to secure false teeth formed out of ivory and bone. However, the modern era of using metallic surgical implants to fix damaged bones only began at the beginning of the twentieth century. Of the metals and alloys available at that time, vanadium

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steel, developed in 1912, offered the best combination of corrosion resistance and mechanical strength. Although this alloy had poor tissue compatibility, it was not until the 1930s that austenitic stainless steels became sufficiently available as a viable alternative, first with compositions similar to that of grade 304 and later with molybdenum additions, that is, evolving towards grade 316L, which was endorsed by the American College of Surgeons in 1946.1 In 1981, a 22Cr13Ni5Mo stainless steel was introduced for hip implants, but this was rapidly superseded by the introduction of high-nitrogen austenitic stainless steel, which offered the same corrosion resistance but without the need for a high-nickel content to maintain the FCC structure, dramatically reducing costs. In response to fears of nickel toxicity, the beginning of this millennium saw the introduction of nickelfree or low-nickel austenitic stainless steels implants, such as the nitrogen-strengthened 23Mn21Cr1Mo (ASTM F2229-02). Cobalt–chromium–molybdenum alloys were first used in dentistry in the late 1920s and then for surgical implants in the late 1930s. The high corrosion resistance and apparent good biocompatibility made these alloys a favored material for orthopedic devices, but during the 1950s, the better strength-to-weight ratio of titanium meant that it started to increase in popularity; CoCrMo alloys remained the preferred choice in the United States up to the late 1960s. In 1974, the Ti6Al4V alloy,

Table 1

which has a much higher ultimate tensile strength than the commercial purity titanium, was introduced for trauma implants, and by the end of that decade, it had become the material of choice in orthopedic surgery, although CoCrMo alloys were still regularly used. The late 1980s and 1990s saw the beginning of the introduction of advanced materials for biomedical applications, including rare earth magnets and NiTi shape-memory alloys. In the present decade, zirconium alloy joint prostheses were introduced and important advances have been made in the application of wear-resistant titanium nitride coatings.1 Undoubtedly, the nature of surgical and dental implant materials will continue to evolve in the coming decades.

2.27.3 Health Effects Related to Corrosion in Body Fluids There is no doubt that surgical implants do raise the level of metallic elements in the body, as illustrated by the data displayed in Table 1.2–34 However, the form of the metal released into the body is often uncharacterized, for example, particulate matter or soluble ions, with the role of serum protein binding being virtually unknown. Furthermore, there is still no agreement over what constitutes safe levels for metals within human body fluids and tissues. Although cobalt,

Approximate concentrations of metals in the human body with and without a total-joint replacement implant2–4 Metal ion concentrations: body fluids (mM); tissues (ppm)

Serum Blood Liver Lung Spleen Lymphatic Heart

Normal Implant Normal Implant Normal Implant Normal Implant Normal Implant Normal Implant Normal Implant

Co

Cr

Mo

Ni

Ti

Al

V

0.003 0.007 0.002 0.33 120 15 200 – – 30 16 000 10 390 30 280

0.001 0.006 0.058 2.1 <14 1130 – – 10 180 690 690 30 90

– – 0.009 0.104 – – – – – – – – – –

0.007 <0.16 0.078 0.50 – – – – – – – – – –

0.06 0.09 0.35 1.4 100 560 710 980 70 1280 – – – –

0.08 0.09 0.48 8.1 890 680 9830 8740 800 1070 – – – –

<0.02 0.03 0.12 0.45 14 22 26 23 <9 12 – – – –

Source: Hallab, N. J.; Jacobs, J. J.; Gilbert, J. L. In Joint Replacement and Bone Resorption; Shanbhag, A., Rubash, H. E., Jacobs, J. J., Eds.; Taylor & Francis: New York, 2006; pp 211–254.

Corrosion in Body Fluids

chromium, and nickel are essential trace elements, in excess, all these lead to carcinogenesis as well as other ailments;2 likewise, vanadium leads to cardiac dysfunction and hypotension,2 while iron has been linked with Parkinson’s disease.5 Excess aluminum can cause anaemia,2 but claims that it causes Alzheimer’s disease have been proved to be completely unfounded.6 Titanium is a nonessential element, which is usually considered inert, but pulmonary disease has been reported in titanium production workers as well as in rats exposed to TiO2 dust.2 The potential to cause cancer is obviously a major concern with metallic implants. Memoli et al.7 reported a slight increase in sarcoma cancers in rats when implanted with devices with high Cr, Co, or Ni contents, and a significant number of implant site tumors have been reported in cats and dogs with stainless steel implants, but fortunately to date, malignant tumors associated with implants in humans are rare.8 Nevertheless, the number of reported cases is increasing and is likely to continue to do so as the patient age decreases and the life-expectancy increases.9 Health concerns with orthodontic devices are also mainly associated with metal release, particularly nickel from stainless steels and the NiTi shapememory alloy.10 With respect to amalgams, apart from the release of toxic mercury, the main concern is the development of lesions of the oral mucosa that can lead to leukoplakia, lichen planus, and oral cancers. Early reports linked these lesions to the potential associated with galvanic cells that can exist between amalgams and precious metal restoration, and hence these have been termed galvanic lesions.11,12 However, galvanic cells appear to be commonplace among healthy populations; Phillips et al. 13 did not find any evidence of leukoplakia in rats that were subjected to galvanic currents.14 More recent studies have found that the lesions are usually caused by contact hypersensitivity to mercury15; the oral mucosa may be either in direct contact with an amalgam or contain some mercury deposits resulting from amalgam corrosion.16 Clinical signs of amalgam corrosion have been reported to be significantly more frequent in patients suffering from oral lichen planus than in controlled groups, and of course, any galvanic cells between amalgams and precious metals would accelerate the corrosion and thus exasperate the problem.17 Nevertheless, lesions of the oral mucosa caused by amalgam restorations are rare and can usually be solved by replacement of the amalgam.15

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2.27.4 Environments Encountered in Biomedical Applications Upon detecting the presence of a foreign entity, the body will produce enzymes and attempt to ingest the foreign object. If the foreign object is too large to be ingested, such as a surgical implant, the body will isolate and encapsulate the object in a fibrous tissue membrane. As such, the implant remains continuously exposed to extracellular tissue fluid. Although the actual compositions of body fluids are complicated, in terms of corrosiveness, the most important characteristics are the chloride, dissolved oxygen, and pH levels. The biological components in body fluids, for example, phosphates, cholesterols, and phospholipids, are usually considered either to play no role in the corrosion process or to exist at insignificant levels. Figure 1(a) shows the typical environmental conditions expected within a range of different body fluids, superimposed on the Pourbaix diagram for chromium in the presence of chloride ions.18,19 From this diagram, it can be predicted that stainless steels and cobalt– chromium–molybdenum alloys are likely to suffer corrosion in many of the environments found within the body, but titanium would be in the passive state for virtually all physiological solutions (Figure 1(b)). Fortunately, the body fluid most likely to be encountered by an implant is blood, which contains 0.9% NaCl and under normal conditions is at pH 7.4, with a redox potential in the vicinity of 0.0 V versus standard hydrogen electrode (SHE) and body temperature of 37  C. Under these conditions, many stainless steels and cobalt–chromium–molybdenum alloys can be expected to be in the passive state. A review by Solar20 in 1979 concluded that inorganic solutions based on 0.9% NaCl were satisfactory substitutes for human body fluids when studying the behavior of passive metals. As a result, the majority of in vitro corrosion experiments have been conducted in either 0.9% NaCl or standard isotonic solutions, such as SBF-K9, Ringer’s, or Hank’s solution, in which additions of bicarbonate and calcium chloride tend to be the main difference to a simple NaCl solution. Dissolved oxygen levels in blood are lower than in saline solutions exposed to air atmospheres, by factors of about 2 and 6 for arterial blood and for veinal blood, respectively. Conversely, bicarbonate levels are about 20 times higher in blood than in saline solutions (Table 2).21–23 Usually, no attempt is made to lower the dissolved oxygen content of the isotonic NaCl solutions to that of the veinal blood, which may

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–2

0

2

4

6

8

10

12

14

2

Potential (V vs SHE)

1.6

Corrosion b

1.2

1.2

0.8

Interstitial fluid Intercellular fluid

Saliva

0.4

Gastric fluid

a

0

0.8 0.4 0

–0.4 Passivation

Urine bile –0.8

–0.4 –0.8

–1.2

–1.2 Immunity

–1.6 –2

0

2

4

6

–1.6

8

10

12

14

16

8

10

12

14

16

pH

(a)

–2

0

2

4

6

2

Corrosion

1.6 1.2 Potential (V vs SHE)

Table 2 Comparison of oxygen and carbon dioxide levels in real and simulated human body fluids.21–23

2

1.6

1.2 Interstitial fluid Intercellular fluid 0.8

Saliva Gastric Passivation fluid

0.4

0

0.4

a

0

–0.4

–0.4 Urine bile

–0.8

–0.8

Corrosion

–1.2 –1.6

–1.6 –2 –2.4 –2

2 1.6

b

0.8

–1.2

(b)

16

–2

Immunity 0

2

4

6

8

10

12

14

–2.4 16

pH

Figure 1 Representative environmental conditions for various body fluids superimposed on Pourbaix diagrams for (a) chromium in solutions containing chloride and (b) titanium. The shaded zones represent the conditions for physiological solutions as suggested by Schenk.19 Adapted from Blackwood, D.J.; Seah, K.H.W.; Teoh, S.H. In Engineering Materials for Biomedical Applications; Teoh, S.H., Ed.; World Scientific: Singapore, 2004; pp 3.1–3.56, with permission from World Scientific.

explain some of the minor differences between the in vitro and in vivo corrosion behaviors of surgical implant materials.22,24 In addition, the minor components in blood, which are usually ignored in in vitro studies, have occasionally been blamed for accelerated

Human artery blood Human veinal blood Rabbit artery blood Rabbit veinal blood NaCl solution

PO2 (mmHg)

PCO2 (mmHg)

HCO 3 (mM)

85–100

35–45

25

40

42–48

25

78

31

20

28

39

23

160

2

–

in vivo corrosion rates, for example, it has been reported that proteins can lower the pitting potential of 304L stainless steel.25 Overall, it appears that typical body fluids are slightly less aggressive than is seawater, indeed Zitter26 recommends that if a stainless steel is to be used to fabricate a biomedical device, it should have a pitting resistance equivalent number (PREN) not less than 26, somewhat below the value of 40 usually required for stagnant seawater. Finally, it needs to be understood that the surgical operation plus the presence of the implant itself may cause the surrounding tissue to undergo severe pathological changes that can result in the development of a more corrosive environment. For example, Laing27 reported that the buildup of hematomas, a condition that could last several weeks, can force the pH in the vicinity of a freshly inserted surgical implant down to as low as pH 4.0. Likewise, the initial stages of the inflammatory response can lead to the generation of hydrogen peroxide and elevated protein levels, both of which may favor the initiation of pitting and crevice corrosion.21,28,29 Furthermore, the degree to which pathological changes occur depends not only on the biological activity of any released corrosion products, but also on the size and shape of the implant, which means that the extent of the pathological changes will vary across the surface of an implant. This could lead to the development of electrochemical cells and the potential gradients necessary to drive localized corrosion. The environment within the oral cavity is not well defined. There are a number of recipes for artificial saliva, of which the most common is that of Fusayama30 (NaCl, 0.400 g dm3; KCl, 0.400 g dm3; CaCl2H2O, 0.795 g dm3; NaH2PO4H2O, 0.69 g dm3; Na2S9H2O, 0.005 g dm3; pH 5.5), which may be acidified with citric acid to pH 4.0. However, in reality, the makeup of human saliva varies considerably between individuals, especially in the sulfide

Corrosion in Body Fluids

content, which can cause tarnishing of both silverand gold-based amalgams. In any case, many foodstuffs are acidic with high-chloride levels, and are thus far more corrosive than saliva. Moreover, oral hygiene has a strong affect on the corrosiveness of the oral environment. Many dental products also contain fluoride, some of the specialist varnishes used by dentists being over 2 wt% fluoride, which can cause corrosion even in commercial purity titanium.31 An example of the problem of defining the environment for the oral cavity is illustrated in the work of Schiff et al.,32 who investigated the corrosion performance of different orthodontic wires. All showed excellent resistance in artificial saliva, but in monofluorophosphate-containing mouthwashes, NiTi-based alloys were found to suffer strong corrosion, while TiNb alloys and TiMo-based alloys gave satisfactory performance. However, when a mouthwash containing stannous fluoride was used, the NiTi alloys outperformed the other two alloys, the TiMo alloys corroding badly.

2.27.5 Metals and Alloys Used in Biomedical Applications In addition to an alloy’s mechanical and corrosion resistance properties, its surface finish is also an important aspect for biomedical devices. The type of surface finish required on an implant is very much dependent on its final use. For example, on the one hand, for prostheses implants usually a rough surface is desirable, as this promotes attachment to the surrounding bone or cement. On the other hand, implants that come into contact with the blood stream, such as vials or reservoirs used to hold the chemicals required in chemotherapy, need to have smooth surfaces to prevent provoking thrombus or crystal formation of the applied chemotherapeutics. As a result, the compositions, mechanical properties, and surface finish of the major types of metals and alloys available are governed by a range of international and national standards, most notably, the various parts of ISO 5832. 2.27.5.1

Titanium and Titanium Alloys

Titanium has excellent corrosion resistance to most environments likely to be found in vivo, with the possible exception of acid anoxic regions (e.g., gastric fluids) where the protective passive oxide may not form. The ability of the passive oxide film to provide corrosion

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protection can be improved by anodizing; note that earlier concerns that anodizing may reduce titanium’s resistance to stress corrosion cracking (SCC) and corrosion fatigue appear to be unfounded.33 Titanium alloys have even better strengthto-weight ratios than does pure titanium, but not quite as high a resistance to pitting corrosion as the parent metal has, and problems can be encountered if the local redox potential is high and the pH low. Theoretically, these conditions could exist during the initial stages of the inflammatory response following surgery, when both a reduction in pH and hydrogen peroxide production can occur, but the author is not aware of any such clinical cases.21,27–29 Overall, titanium alloys have better corrosion resistance than do cobalt–chromium–molybdenum alloys and stainless steels. In the previous decade, concerns about the cytotoxicity of vanadium have led to the development of a number of new alloys as potential replacements for the dual phase (a þ b phases) Ti6Al4V alloy; there have been reports of elevated metal levels in soft tissues and bones surrounding implants fabricated from this alloy.34 Many of the potential replacements are b stabilized, which allows a lower elastic modulus and thus should aid bone growth, but at the cost of lower wear and fatigue resistances. However, two replacement alloys, Ti15Mo5Zr3Al and Ti6Al2Nb1Ta0.8Mo, have already been used in hip prostheses in cemented and noncemented applications, respectively, with apparently excellent clinical results.35 Despite their excellent corrosion resistance and biocompatibility, titanium and its alloys are still not the perfect biomedical materials, as their poor shear strength makes them unsuitable for screws and other forms of fastener devices. This can lead to fixation problems if galvanic corrosion is to be avoided between the titanium alloys and attaching screws. In addition, titanium alloys also have a high coefficient of friction, which means that wear particles may form if rubbing against bone or another implant surface occurs; the latter case has led to failures due to fretting corrosion, in which the passive oxide film is worn away.36 The poor fretting resistance of the Ti6Al4V alloy represents its most serious drawback for use in load-bearing prosthesis, and therefore, considerable effort has been made to find possible solutions.37–40 Encouraging results have been reported for both anodizing and titanium nitride coatings.37,41 Anodization has the advantages of low cost and ease of operation, while the nitride-coated Ti6Al4V has the better fretting resistance38; both techniques have demonstrated

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biocompatibility.37,42 Recently, a zirconium alloy containing 2.5% niobium has been introduced for joint prosthesis. A thick oxide layer formed on the Zr2.5Nb by heating to 535  C gives it a far superior wear resistance to that of titanium alloys.43 Zr alloys have similar corrosion resistance properties to titanium, but are difficult to machine as they show pyrotechnic tendencies. 2.27.5.2 Alloys

Cobalt–Chromium–Molybdenum

The main alloys used are based on either CoCrMo alloy, which has been used extensively in dentistry and more recently for artificial joints, or CoCrNiMo, which has a very high ultimate tensile strength and thus is used for making the stems of prostheses for heavily loaded joints. Molybdenum is added to improve the mechanical properties by decreasing the grain size, rather than to improve the corrosion resistance. These alloys have excellent resistance to most forms of corrosion, including crevice corrosion and corrosion fatigue, but fretting corrosion can cause failures. The nickel-containing alloys have the better corrosion resistance, and ASTM F 1058 (40Co20Cr15Ni7Mo) has a long track record as a permanent implant alloy.44 However, concerns about the release of toxic Ni2+ ions have resulted in the CoCrMo alloy, ASTM F 75 being the dominant cobalt–chromium–molybdenum alloy in use today. A final concern with using CoCrMo alloys is the potential release of chromate, a known carcinogen, into the body. This worry also applies to stainless steels, since these too contain chromium, although at lower levels. 2.27.5.3

Stainless Steels

Early attempts to use 12% Cr Sherman plates with type 304 stainless steel screws led to predictable galvanic corrosion problems; nonetheless, some of these plates remained in patients for over 30 years, only being removed when tumors (nonmalignant) develop over the corroding implants.45 The galvanic corrosion problem was solved by using an all type 304 stainless steel construction, but this did not prevent problems associated with pitting and crevice corrosion. Nevertheless, small type 304 stainless steel Sherman plates have been removed from patients after 30 years of service without any sign of corrosion; possibly the low oxygen content of body fluids prevented the pitting potential from being exceeded.21 In contrast, a

number of cases of larger type 304 stainless steel implants developing localized corrosion problems shortly after implantation have been reported.46 These apparently contradictory results bear testament to the importance of the pathological changes that accompany the healing process in the first few weeks after surgery, and the extent of the healing process can be expected to be dependent on the size of the implant.21 The development of the molybdenum-containing type 316L stainless steel led to a significant decrease in the number of failures related to localized corrosion. However, from a review of failures during the years 1980–1989, Zitter26 suggested that a PREN greater than 26 was required to prevent in vivo pitting corrosion, which is slightly above the PREN of most type 316L stainless steels produced, making this grade more suitable for temporary implant devices. The PREN value can be pushed above the recommended threshold of 26 by the addition of nitrogen, which also has the advantage of increasing the ultimate tensile strength of the material, but at the expense of a lower elongation at fracture. During the last two decades, the importance of sulfide and phosphide inclusions in pit initiation processes has been recognized and this has led to the development of type 316LVM stainless steel, in which the nonmetallic inclusion content is reduced by vacuum melting. The composition of the 316LVM grade is usually slightly above the normal 316 specification, typically being 18Cr14Ni3Mo, such that its PREN value of 28 is above Zitter’s threshold limit. The recently developed low-nickel stainless steels such as 23Mn21Cr1Mo (ASTM F 2229-02) are nitrogen strengthened, which can also push their PREN beyond 26. The superior mechanical and formability properties of stainless steels over titanium alloys are a great advantage in orthodontic applications. However, the consumption of acidic food and beverages means that there is a much greater risk of localized corrosion in the oral cavity than inside the body. To counter this threat, ultraclean high-nitrogen austenitic stainless steels have been developed, for example, 21Cr10Ni3Mo0.3Nb0.4N.47 2.27.5.4

Nickel–Titanium Alloy

NiTi is a shape-memory alloy with super-elasticity, a property that is of interest in both surgical and dental applications. Mantovani48 reviewed the possible uses of NiTi in medical appliances, typical applications

Corrosion in Body Fluids

including dental braces, medical staples, and nails. The most likely form of corrosion on NiTi is pitting; its resistance to pitting in body fluids appears to be similar to that of 316L stainless steel type.49 However, NiTi fares worse than stainless steel orthodontic wires in acidified artificial saliva.50 Given that there is some evidence that nickel initiates cancer, the high nickel content of NiTi is a matter of concern.2 To date, studies indicate that the amount of nickel released into the body from the corrosion of NiTi depends on the local environment and the tests in artificial saliva showed similar Ni release rates to 316L stainless steel,51 while in simulated body fluids, the NiTi released three times as much nickel as in artificial saliva.52 Rather alarmingly, Heintz et al.53 found that when fibroblast cells adhere to the NiTi alloy, they damage the oxide film, leading to rapid localized corrosion and that this has caused failures in stent wires of explanted endovascular grafts. Hashimoto and Morita54 have also reported significantly higher nickel ion release rates in the presence of living fibroblast cells. The mechanism by which this accelerated corrosion occurs is not yet known, possibly the attached fibroblasts lead to the development of occluded electrochemical cells. Nevertheless, the biocompatibility of NiTi, as determined in shortterm in vivo tests on guinea pigs, was reported to be comparable with type 316LVM stainless steel.55 This view appears to be supported by the largely favorable surgical clinical evidence available to date, including excellent biocompatibility in tendon tissue, minimal corrosion on retrieved implants, and nickel concentrations in the major organs at levels similar to those for stainless steel devices.56 However, the concerns about potential Ni release and its poor performance in dental applications have led to a number of efforts aimed at improving the NiTi alloys’ resistance to pitting corrosion, the most promising techniques being diamond-like carbon coatings and nitriding.57–59 Recently, a new nickel-free shape-memory alloy has been developed, Ti18Nb4Sn, with a corrosion resistance in simulated body fluids apparently comparable with that of commercial titanium.60 Once acceptable biocompatibility of this alloy has been demonstrated, there is a good chance of it eventually replacing NiTi, at least in orthodontics. 2.27.5.5 Porous Materials and Metallic Foams As demonstrated earlier, titanium alloys and CoCrMo alloys have excellent corrosion characteristics for the

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construction of surgical implants. However, the elastic moduli of these solid metallic alloys are much higher than those of human bone, which means that stresses are not transferred to the surrounding bone effectively, leading to irregular bone growth and osteoporosis. One possible solution to this problem is to use porous implant materials that have lower elastic moduli closer to those of human bone. Such porous materials not only have the advantage of a more suitable elastic modulus, but also open up the possibility of allowing bone growth into the implant itself, thereby improving the adhesion and thus reducing the likelihood of fatigue or fretting failures. Unfortunately, the corrosion rates of porous titanium, porous CoCrMo, and porous NiTi in simulated body fluids have all been reported as being significantly higher than those of their solid counterparts, possibly due to crevice corrosion within the pore matrix.61–64 Very recently, interest has been shown in metallic foams, in which a foaming agent helps to control the porosity that can be up to 80%.65 It has been reported that, for such titanium foams, corrosion rate as low as 0.07 mm year1 can be obtained after aging for 6 days in a simulated body fluid (0.01 M phosphate buffer pH 7.4, 0.027M KCl þ 0.137 M NaCl), which is comparable with or even better than that of solid titanium.66 However, since the wetted surface area of a foam is more than 1000 times its geometric area, the concentration of metallic ions released is still much higher from foams than from solid materials, increasing the risk of health problems related to elevated metal levels in body tissues. 2.27.5.6

Magnesium Alloys

At first sight, most readers will be surprised at the suggestion of fabricating biomedical devices out of a material that corrodes as readily as magnesium; however, corrosion can in fact be a desirable feature! Preliminary investigations are being conducting into the use of magnesium alloys as degradable implants for musculoskeletal surgery,67 the idea being that magnesium ions encourage bone cell activation so that the bone slowly regenerates as the metal corrodes. This represents an interesting new challenge to corrosion scientist: how to tailor the corrosion rate of the Mg alloy to match the bone regrowth rate? However, it should be noted that Witte et al.67 reported that in vivo corrosion rates were four orders of magnitude lower than in vitro rates, and that the order of relative performance between the various Mg alloys tested changed with environment, leading these

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authors to conclude that the present ASTM standard in vitro corrosion tests are unable to predict the in vivo corrosion rates of magnesium alloys. 2.27.5.7

Rare Earth Magnets

There are a number of ternary alloys containing rare earth elements that have remarkably strong magnetic properties, such as the samarium–cobalt family. Although originally developed for the magnetic storage industry, these alloys are used in a number of specialized medical applications, for example, as dental keepers; the strong magnetization is used to keep dental fixtures in place. Unfortunately, these rare earth magnets have very poor corrosion resistances and cannot be directly inserted into any body fluid. One solution is to completely seal the magnet inside a stainless steel cladding, but this must not reduce the effectiveness of the magnetization, which rules out the austenitic steels. In the case of dental keepers, ferritic stainless steels with chromium levels as high as 55% have been used, sufficient to withstand the most corrosive foodstuffs. The magnets hold the dental fixture to a ferromagnetic material usually cemented into a residual tooth root; typically this is a Pd–Co alloy offering good corrosion resistance, but there are concerns about the leaching of cobalt ions, which are cytotoxic at high concentrations.68 2.27.5.8

critical. Furthermore, it has been shown that for most people, the major route for mercury to be taken into the body is via food, less than 10% coming from dental amalgams.73

Dental Amalgams

Dental amalgams are high-strength multiphase alloys, which makes them vulnerable to localized galvanic or intergranular corrosion between the different phases. The majority of modern dental amalgams are prepared from two types of alloys: conventional silver tin amalgam and high-copper amalgams. The high-copper amalgams have superior clinical properties with a higher resistance to corrosion.69 The corrosion of any amalgam is of concern, as it leads to the release of mercury into the body, and in rare cases, can cause oral lesions if this redeposits in the oral mucosa.15 In conventional silver tin amalgams, the most base phase is g2 (Sn7Hg), which releases mercury when it corrodes.70 Conversely, the most corrosion-prone phase in high-copper amalgams is Z0 (Cu6Sn5), preferential corrosion of which does not release mercury into the body and thus these alloys have recently been favored.71 However, Joska et al.72 found that mercury release rates from conventional silver amalgams and high-copper amalgams were very similar, the method of preparation being

2.27.5.9

Titanium Nitride Coatings

Very hard yet smooth, low friction coatings of titanium nitride, typically a few microns thick, can now be produced by a variety of methods, including ion implantation,38 nitriding,39,40 physical vapor deposition,41 and magnetron sputtering.74 These coatings have been shown to have excellent biocompatibility and have been successfully used on Ti6Al4V to protect against fretting corrosion and other forms of wear, allowing the world’s first collar bone replacement to be preformed in 2003,41 as well as on NiTi alloy and stainless steels to provide protection against pitting and crevice corrosion.59,75 The excellent combination of biocompatibility with corrosion and wear resistance means that TiN coatings will undoubtedly find increased usage in biomedical applications. One potential concern about nitride coatings that has yet to be fully investigated is the possibility of galvanic corrosion of the underlying metal at holidays or damaged areas of the coating. This has already been shown to be a problem at microstructural defects in TiN coatings on type 304 stainless steel in saline solutions.76 Variations on titanium nitride coatings are also being developed, such as titanium aluminum nitride, that reportedly provide better protection to stainless steels without any loss of biocompatibility.76,77

2.27.6 Corrosion Types Encountered in Biomedical Applications 2.27.6.1

General Corrosion

For a successful implant material, the longterm general corrosion rate should certainly fall to much less than 1mm year1; for almost any other application, such low corrosion rates would be considered insignificant. Nevertheless, even at these rates, it has been reported that after implantation the nickel, chromium, and cobalt levels in surrounding tissues can be significantly higher than normal values (Table 1).2 It has also been shown that the presence of metallic ions released from nonmolybdenum-containing highnitrogen stainless steels suppresses cell growth of human gingival fibroblasts.78

Corrosion in Body Fluids

2.27.6.2

Pitting Corrosion

Although cases of pitting corrosion were common with the early stainless steel implants fabricated from grade 304, the addition of Mo (2–3%) to form 316L grade stainless steel has greatly reduced the number of failures. Zitter26 has suggested that a PREN greater than 26 is required to prevent in vivo pitting corrosion. Stainless steel biomedical devices should thus be manufactured from at least high-quality 316L, preferably from ultraclean grades such as 316LVM or grades containing nitrogen additions. The risk of pitting corrosion of stainless steels in the oral cavity is higher than that in implants because of the number of chloride-containing acidic foodstuffs regularly introduced into the mouth. Although this is partly compensated by the ease by which orthodontic devices can be retrieved, it is still recommended that ultraclean highnitrogen austenitic stainless steels be used rather than the standard 316L grade. One of the major concerns about the use of surgical implants based on cobalt–chromium– molybdenum alloys is that pitting corrosion could lead to carcinogens being released into the body. This has resulted in numerous in vitro investigations into the pitting behavior of these alloys in pseudobody fluids, all of which have reported excellent resistance to pitting as long as static conditions are maintained.64,79 However, pitting corrosion has been observed when the CoCrMo alloys were subjected to either cyclic loads or severe cold working.80 Titanium metal is immune to pitting corrosion in any in vivo environment likely to be encountered. Although Ti alloys are less resistant, no in vivo pitting-related failures have been reported. The pitting behavior of nickel–titanium shape-memory alloys has already been discussed earlier.

2.27.6.3

Crevice Corrosion

The most common location of crevice corrosion found in biomedical applications is beneath the heads of fixing screws and it is a very serious problem with stainless steel devices, even when fabricated from high-grade molybdenum and nitrogen-containing alloys. In a survey conducted in 1959, Scales et al.81 found that 24% of grade 316 stainless steel bone plates and screws removed from patients revealed signs of crevice corrosion. Although the introduction of the low nonmetallic inclusion type 316LVM stainless steel and the use of an austenitic microstructure free of any d-ferrite phase have reduced the level of

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crevice corrosion problems, these have not been eliminated. Crevice corrosion on CoCrMo alloys appears to be less of a problem than on stainless steels. Syrett and Davis82 found no crevice corrosion on specimens removed from dogs and rhesus monkeys after 2 years of implantation. Similarly, Galante and Rostoker83 found no crevice corrosion on CoCrMo alloy implants removed from rabbits after 1 year, although the latter authors did find single pits in the crevice regions that might have eventually developed into crevice corrosion if given sufficient time. Crevice corrosion of titanium in neutral chloride environments has only been reported at temperatures in excess of 70  C, that is, it is not expected in vivo. As with pitting corrosion, titanium alloys are less resistant to crevice corrosion than is pure titanium, and Galante and Rostoker83 have reported single pits in the crevice regions of Ti6Al4V specimens implanted in rabbits for 1 year, but no actual crevice corrosion. Despite titanium’s excellent resistance to crevice corrosion, it is not the answer to the problem, since its poor shear strength makes it unsuitable for screws and other fasteners, which are the main crevice formers. Finally, some total-joint implant designs contain metal-on-metal press-fit conical tapers, which are subjected to stress and motion. Retrieval studies have found crevice corrosion at taper connections consisting of CoCrMo alloys heads with both CoCrMo alloy and Ti6Al4V stems.2,84 Occasionally, titanium alloy stems have been attacked, which suggests that the wearing away of the passive film plays a role, that is, the mechanism is a combination of fretting and crevice corrosion.36 2.27.6.4

SCC and Hydrogen Embrittlement

To the best of the author’s knowledge, neither SCC nor hydrogen embrittlement has been observed on recovered surgical implants. Although retrieved implants may show evidence of cracks, they do not show the physical characteristics associated with SCC, and thus almost certainly, result from mechanical damage either during manufacturing or during the recovery process. The laboratory experiments conducted to date support the presumption that the three classes of common implant alloys (Ti, CoCrMo alloys, and stainless steels) are not susceptible to SCC in in vivo environments. The only counter evidence comes from in vitro tests conducted under experimental conditions that are highly unlikely to ever exist in any true in vivo situation, such as extreme

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negative potentials or acidic MgCl2 solutions. However, Rodrigues et al. have very recently reported the first evidence of hydrogen absorption being implicated in the corrosion of retrieved titanium alloy hip implants.106 2.27.6.5

Corrosion Fatigue

One of the features that distinguish corrosion fatigue from mechanical fatigue is that the former is strongly dependent on the frequency of the applied loaded, with low-frequency cycles causing the most damage. Unfortunately, many medical devices are subjected to low-frequency loads, for example, simply walking represents a cyclic loading at about 1 Hz on a hip implant, so the threat of corrosion fatigue might be expected to be high. However, in 1975, Bechtol85 reviewed all types of clinical fatigue-related failures and claimed that the most common root cause of failure was not related to corrosion, but rather to a breakdown of the bone–cement support interface. This leads to a widening of the separation between the metal prosthesis and bone–cement and finally to the deformity of the metal stem. This mechanism was recently supported by von Knock et al.,86 who found no evidence of corrosion on 11 CoCrMo alloy femoral components retrieved after 2–15 years of service and suggested that the majority of the micromotion between the prosthesis and bone occurs at the bone– cement/bone interface. Likewise, reviews of the literature related directly to corrosion fatigue of prostheses implants by Leclerc87 in 1982 and Zitter26 in 1991 both concluded that as long as the manufacture and metallurgical condition of the device conformed to international standards (e.g., ISO 5832 or ASTM F 138), corrosion played only a minor role in most fatigue failures. However, in contrast, Morita et al.22 reported that the fatigue strengths of 316 stainless steel and a CoCrNiFe alloy were considerably less in vivo (rabbits) than in air and proposed that this was due to the low dissolved oxygen concentration in body fluids, causing a corrosive action on the alloys. Furthermore, in his 1982 review, Leclerc did note that the longer the prosthesis was implanted in the patient, the greater the role of corrosion, which is significant giving that, as explained in the introduction, it is anticipated that the required service life of surgical prosthesis will increase in the coming decades. If corrosion fatigue does become an important issue in the future, the importance of material selection and design were emphasized by Piehler et al.,88

who tested hip nail plates and found that large plates had better corrosion fatigue resistance than did small ones and that Ti6Al4V outperformed 316L stainless steel. The good performance of titanium alloys was also highlighted by Hughes et al.,24 who reported that the corrosion fatigue resistance of titanium was virtually independent of pH over the range 2–7, whereas that of stainless steel declines rapidly below pH 4. These observations are consistent with the findings of Yu et al.89 that corrosion fatigue in stainless steels can be initiated by pitting corrosion. The latter authors also reported that the corrosion fatigue resistance of the common implant alloy Ti6Al4V can be enhanced by nitrogen implantation and heat treatments to produce fine prior-b grain sizes. 2.27.6.6

Fretting Corrosion and Wear

Because all the successful surgical implant alloys are based on passive metals, any process that wears away the protective oxide film is of major concern. As a result, fretting corrosion represents the most important form of attack on load-bearing prosthesis and all three major classes of alloys used, namely, Ti alloys, CoCrMo alloys, and stainless steels, suffer fretting corrosion, sometimes in combination with crevice corrosion.90 Fretting corrosion not only results in metal loss but also alters the dimensions of the prosthesis, causing fixation problems and allowing additional micromotions. This in turn can increase mechanical wear and lead to the loss of the surrounding bone– cement or bone, which, besides being a serious problem in itself, increases further the amount of movement of the implant, thereby increasing the likelihood of fatigue-related failures.85 The situation is made worse by the fact that the corrosion products collect locally as particles that can cause further abrasion of the implant; for example, black titanium oxide debris is often found in the vicinity of implants.91 The main cause of the shearing micromovements that eventually lead to the fretting corrosion is believed to be the large difference between the elastic moduli of solid metallic implants and the surrounding bone or bone–cement. Morita et al.92 investigated the wear resistances of a high-nickel version of 316L stainless steel (Fe17Cr14Ni2Mo), a cobalt–chromium–molybdenum alloy (Co28Cr6Mo), and the titanium alloy Ti6Al4V and found that rubbing between metal and ultra-highmolecular weight polyethylene, commonly used as a liner for the sockets into which the ball of hip implants are inserted, only caused the oxide on the Ti6Al4V alloy to suffer damage, that is, the titanium alloy was

Corrosion in Body Fluids

the least resistant. Worse, when metal rubbed metal, the oxides on all the three implant alloys were destroyed, even at low loads. Furthermore, the addition of calcium chloride and/or hydrogen peroxide to saline solutions has also been shown to lead to increased fretting corrosion of Ti6Al4V.93 Since H2O2 can be produced as part of the inflammatory response of damaged body tissues, it can be postulated that new load bearing prosthesis are at the most risk of developing fretting corrosion, although the corrosion may subside as the H2O2 production ceases, sufficient loosening of the device might occur to induce mechanical fatigue. Efforts to reduce the effects of fretting are based mainly on coatings, usually titanium nitride or aluminum oxide.38–41,74,94 Besides wear-resistant coatings, endeavors have been made to reduce the threat of fretting corrosion by improving the binding between the implant and its surroundings, be it bone or bone–cement. Proposals of how to achieve this include engineering the shape, topography, porosity, or composition of the implant to provide either in-growth of tissue or enhanced ongrowth of mineralized bone61,95; plasma spraying a titanium coating with a specific surface roughness on the surface of the Ti6Al4V96; or depositing strongly adhered hydroxyapatite coatings that can fuse with the growing bone. Encouraging results have recently been obtained for silicon-doped hydroxyapatite coatings.97 Although some reports have suggested that hydroxyapatite coatings do not provide any longterm improvement in fixation,98 overall clinical results indicate that coated implants perform well, especially in young patients.2,99 Occasionally, erosion– corrosion problems are encountered on implanted valves and pumps, as with nonbiomedical application, the solution to this problem is to use a more resistant material, such as titanium.100 2.27.6.7

Galvanic Corrosion

Galvanic corrosion has certainly caused the failure of a number of biomedical devices. However, the vast majority of these cases were caused by poor quality control or a lack of appreciation of the existence of galvanic couple between apparently similar materials. There have even been examples of galvanic corrosion-related failures arising, because just a single grade 304L screw was used in what was otherwise an all grade 316L construction101; since it is not possible to visually distinguish one grade of stainless steel from another, the solution to the problem is careful quality control. When the correct materials

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have been used, galvanic corrosion is not normally a problem; however, the poor shear strength of titanium alloys means that it is not suitable for use as fasteners, so there may be times when the production of a galvanic couple is unavoidable. In the event that titanium and any cobalt–chromium–molybdenum alloys or stainless steels are coupled together, it is likely that the former will become the cathode and thus accelerated corrosion of the latter alloys may be anticipated. However, titanium and its alloys are easily polarized and their passive films make them poor cathodes, which in practice means that the extent of accelerated corrosion caused to any metal from coupling to a titanium alloy can be expected to be small. This argument has been confirmed in a literature review by Mears102 and also for titanium/cobalt– chromium–molybdenum alloy combinations by the in vitro experiments of Lucas et al.80 and in clinical use as reported by Jackson-Burrows et al.103 Nevertheless, Rostoker et al.104 found that type 316L stainless steel suffered pitting corrosion in 1% NaCl solution at 37  C when it was coupled to either Ti6Al4V, CoCrMo alloy, or graphite, but no pitting corrosion was found when any two of the other three materials were coupled together. There is also one further aspect of galvanic corrosion to be considered, that is, any bimetallic couple is of course a small battery in which a current flows between the anode and cathode. Even if this current is too small to cause any significant corrosion problems, these could be sufficient to cause the patient pain, indeed persistent pain resulting from such situations has led to the need to retrieve some biomedical devices.105 In dental applications, galvanic corrosion of amalgams has been linked to causing oral lesions, which can develop into cancer, particularly in patients with a hyposensitivity to mercury.15

2.27.7 Conclusions The knowledge of corrosion and mechanical properties of materials has allowed the development of a number of extremely successful biomedical alloys. As a result, as long as the chosen materials match the requirements of national and international standards, the likelihood of a surgical implant suffering a corrosion-related failure is very low. The most important remaining areas of concern are fretting and corrosion fatigue. Even here, recent advances in titanium nitride coatings and fixation techniques are extremely encouraging, suggesting that a solution to

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Liquid Corrosion Environments

at least some of the problems is close at hand. The TiN coatings probably represent the most important advance in the protection of biomedical devices since Leclerc87 wrote his chapter on surgical implants for the third edition of Shrier’s Corrosion in 1994. However, TiN coatings have yet to be demonstrated in the tapered joints of some total-joint replacements where a combination of fretting and crevice corrosion can cause failures even in titanium alloys. Moreover, device failure is not the only concern; now that younger patients are receiving implants, the concern that extended exposure to even very low levels of corrosion products could result in medical complications, including cancer, is becoming more and more important. So there is still a continuing need to further reduce corrosion rates. Furthermore, in recent years advanced materials have been developed that possess properties considered highly desirable for biomedical devices, for instance, porosity, shape-memory, and high magnetism. This trend is likely to continue, as the recent upsurge in funding on life science-related research will produce a number of new materials and devices specifically designed for biomedical applications, such as implanted sensors, automatic drug dispensers, and even micromachines. Techniques to protect these materials without interfering with the very functional properties that make them so desirable have been, and will have to be continually, developed, and occasionally, the biomedical industry will continue to throw out more unusual challenges to the corrosion scientist, as in the case of the sacrificial magnesium alloys being proposed to stimulate bone growth where a controlled corrosion rate is deemed desirable.

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