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Corrosion in titanium dental implants/prostheses--a review.


The general consensus about the most corrosion resistant biocompatible metallic biomaterials are the special metals namely-titanium, niobium, tantalum and their alloys, followed by cobalt based alloys and finally the stainless steel grades [1-3]. The most commonly used implant/ prostheses material used today has been summarized below in Table 1 with their common names, UNS, ASTM, ISO and alloy designations [4-6]. Commercially pure titanium and its alloys are known for their use in medical application owing to their good corrosion resistance, biocompatibility and bioactivity in the human body [7].Titanium and its alloys have been used as prosthetic material in several reconstructive and resective procedures in human body since many years [8]. The use of dental implants in the partial and complete edentulism has become the primary treatment regimen in the modern dentistry [9]. These implants appeared as early as late 1920's and gained widespread usage during the last 2-3 decades.

Although these biomaterials have good mechanical and biological properties there corrosion resistance is still critical for the overall success of the treatment procedure. It has been long recognized that the corrosion products formed as a result of metal-environment interactions have a significant bearing on the biocompatibility and long term stability of the prostheses/implant. The material used must not cause any biological adverse reaction and must retain its form and properties [10-12] during function.

Human stomatognathus is subjected to varying changes in pH and temperature owing to differences in local, systemic, environmental, economic and social conditions for each individual. Corrosion can result from the presence of a number of corrosive species like hydrogen ion ([H.sup.+]), sulfide compounds ([S.sup.2-]), dissolved oxygen, free radicals ([O.sub.2], [O.sup.-]), and chloride ion ([Cl.sup.-]) resulting in the metal surface breakdown and a consequent adverse tissue reactions [13].

Nowadays, osteo-integrated implants/prostheses can be regarded as well proven safe alternatives on a large scale to functionally rehabilitate both partially and completely edentulous patients. One of the chief success factors related to modern titanium implant system is the achievement of a fast and tight interconnection between the implant surface and the bone tissue owing to its low thermal conductivity, low density, lower elastic modulus mismatch compared to bone, high hardness, outstanding biocompatibility and remarkable corrosion resistance [14-16]. The prerequisites of a successful implant system have been summarized in Figure.1.

Physical metallurgy of titanium alloys

Titanium is a transition metal with an incomplete d-shell in its electronic structure that enables it to form solid solutions with most substitutional elements having a size factor within 20% (Hume-Rothery's principles for substitutional and interstitial solutions). In its elemental form titanium has a high melting point (1678[degrees]C), exhibiting a hexagonal close packed crystal structure (a/alpha) up to the beta transus temperature (882.5[degrees]C), transforming to a body centered cubic structure (b/beta) above it [18]. Titanium alloys may be classified accordingly either a, near a or a + b, metastable b or stable b depending upon their room temperature microstriucture. [18-19]. In this regard substitutional alloying elements for titanium fall into three basic categories

1. Alpha stabilizers like Al, O, N, C

2. Beta stabilizers like Mo, V, Nb, Ta (isomorphous), Fe, W, Cr, Si, Ni, Co, Mn, H (eutectoid)

3. Neutral stabilizing elements like Zr.

Alpha and near alpha titanium alloys exhibit superior corrosion resistance but lower strength. On the other hand a + b alloys exhibit higher strength due to presence of both phases in the microstructure. Their properties depend upon composition, the relative proportions of each phase present with their prior thermal treatment and thermo mechanical processing conditions. Beta alloys in contrast have high strength, good formability and hardenability. Beta alloys also offer lower elastic modulus and higher corrosion properties compared to other alpha and mixed alloys there by suggesting lower coefficient of elastic mismatch with the bone and favorable stress distribution and outcome. [20].


A beta alloy is fundamentally defined as an alloy whose chemical composition lies above [b.sub.c] as shown in Figure 2 that is, it contains sufficient amount of beta stabilizing agents to retain 100% body centered cubic microstructure upon quenching from above the beta transus temperature [20-21]. Alloys lying above this critical minimum level of beta stabilizing content may still lie within the two-phase region, making the as-quenched beta phase being metastable with the vulnerability of precipitating a second phase upon aging.

Titanium alloys with increasing alloying content, exceeding a critical beta value are considered stable beta alloys where no precipitation of the second phase takes place during practical long-time thermal exposure. Process variations are traditionally used to control the alloy microstructure and therefore to optimize titanium alloy properties such as ductility, strength, fatigue resistance and fracture toughness. The effects of various microstructures are then correlated with engineering properties and the most common microstructural features studied are the metastable beta alloys, beta grain size, size and age distribution of aged alpha phase. Apart from the alpha phase precipitation transient b' or w phase or several other intermetallic compounds may also be observed in the metastable beta region depending upon alloy composition, heat treatment, processing history and service conditions [22-23].

General Concepts related to corrosion

There are two essential features determining corrosion of an implant

1. Corrosion Thermodynamics [] The application of chemical thermodynamics are primarily related to the oxidation-reduction reactions for the given electrochemical system. The system tends to attain the lowest energy by releasing energy during the reaction [24]. It is represented by the potential (voltage) axis on the Evan's diagram and is given by equation

DG = DG[degrees] + RT lnK for the chemical reaction

DE = DE[degrees] - lnK Nernst Equation for the electrochemical reaction

Where DG, DE are the energy and potential change, DG[degrees] and DE[degrees] are the standard free energy and potential change, R is universal gas constant, T is the temperature under study, n is the number of valence electrons involved in the reaction and K is the equilibrium constant of the given reaction at a specified temperature.

2. Corrosion Kinetics [] The electrochemical kinetics include factors that physically, obstruct corrosion from taking place [24]. These factors tend to hinder the various electrochemical reactions going in the system either through concentration, temperature or velocity kinetics in the system. It is represented by log current density axis on Evan's diagram and is majorly driven by the overpotential equation

[h.sub.conc] = log[1-

Where, [h.sub.conc] is the concentration overpotential, R is universal gas constant, T is the temperature under study, n is the number of valence electrons involved in the electrochemical reaction, [i.sub.c] is the cathodic current density and [i.sub.L] is the limiting current density.

Types of corrosion

Uniform corrosion (Generalized corrosion)

A uniform regular removal of the metal from the surface is usually the most common mode of corrosion. The corrosive environment must have same access to all parts of the metal surface and the metal itself must be metallurgically and compositionally uniform. At times general corrosion in aqueous body fluids like phosphate buffer saline (PBS), ringer's lactate (RL), normal saline (NS) etc may take the mottled form, severely roughened metal surface that resembles localized attack. This uneven localized attack results from variations in the corrosion rate of localized surface patches due to localized masking of metal surfaces by process scales, corrosion products, food lodgment and surrounding and adjacent super structures. When titanium is in the fully passive condition, corrosion rates are typically less than 0.02 mm/yr (0.8 mil/yr) and well below the 0.13 mm/yr (5 mils/yr) maximum corrosion rate commonly accepted for biomaterial design and application. This minimal acceptable corrosion rate is primarily due to the finite +4 oxidation of titanium alloys owing to the formation of adherent Ti[O.sub.2] film although the surface oxide is more complex than a single Ti[O.sub.2] oxide over their surface.

General corrosion becomes a concern at high temperatures in highly acidic environments owing to consumption of hot, spicy and sticky foods. In strong and/ or hot reducing acids (plaque deposits) the oxide film of titanium can deteriorate and dissolve, and the unprotected metal is oxidized to the violet colored soluble trivalent ion ([Ti.sup.3+]) in acid solutions which is further converted to pale yellow [Ti.sup.4+] ion in presence of oxidizing species which on further hydrolysis may form insoluble Ti[O.sub.2] precipitates/ scales and inhibiting subsequent corrosion.

Uniform corrosion for titanium implants can be determined from weight loss data (increase or decrease in weight depending upon the environment and by products in accordance to ASTM G1 & G31), dimensional changes (shape, size, appearance and texture) and electrochemical methods (anodic and cathodic polarization, cyclic voltammetry and electrochemical impedance measurements).

Corrosion rates in millimeters per year for titanium alloys can be calculated from weight loss data as under:

Corrosion rate (mm/yr) =

where d is the titanium implant alloy density (in grams per cubic centimeter which is approximately 4.51g/cc for c.p.Ti), A is the sample surface area (in square centimeters), t is the exposure time (in hours), and W is the weight change (in grams).

Corrosion rates in millimeters per year can be calculated from electrochemical measurements on the other hand using the equation:

Corrosion rate (mm/yr) =

where [i.sub.corr] is the measured corrosion current density (in milliamps per square centimeter), d is titanium alloy density (in grams per cubic centimeter), and EW is the equivalent weight for titanium. The equivalent weight for titanium is approximately 16 under reducing acid conditions and 12 under oxidizing conditions depending upon the number of valence electrons involved. The value of [i.sub.corr] is typically determined from Tafel slope extrapolation or linear polarization methods [25-26]

Pitting corrosion

Localized corrosion attack in an otherwise resistant surface produces pitting corrosion. It usually occurs on base metals which are protected by a naturally forming thin film of an oxide (for instance the firmly adherent Ti[O.sub.2] over the Ti surface) when the potential of the film exceeds the breakdown potential of the oxide in a given environment. In the presence of certain ions like chlorides and sulphides the film locally break downs and rapid dissolution of underlying metal occurs in the form of pits. When the anodic breakdown (pitting) potential of the metal is equal to or less than the corrosion potential under a given set of conditions, spontaneous pitting can be expected.

Because of the protective oxide films the titanium implant surface exhibits anodic pitting potential that are very high ([much greater than] 1V) compared to other biomaterials used (iron, steel, cobalt-chromium alloys etc). Thus pitting corrosion is not of much concern in the oral environment for titanium alloys. For example pitting potentials exceed +5 to +10 V versus the saturated calomel electrode (SCE) in body fluids like chlorides (NS) and lactates(RL) and typically +60 to +80 V in sulfates and phosphates(PBS).Although pitting potential for titanium dental implants is not of much clinical significance but this property of protection potential or repassivation can still be used to define the minimum potential at which pitting can be maintained [27].

Guiding principles under the American Standards for Testing of Materials G3 & G5, electrochemical techniques involving potentiostatic and potentiodynamic testing are used similar to uniform corrosion measurements to study pitting behavior of titanium dental implants [28-29]. Anodic pitting potential requires slow scan rates (d[]0.5 mV/s) and surface condition of the implant surface under test. For example, abraded or sandblasted implants will exhibit significantly lower pitting potentials compared to acid pickled implant surfaces. Repassivation potentials on the other hand can be measured using the galvanostatic technique [27,30] by impressing an anodic current density of approximately 200 mA/[cm.sup.2] on the specimen for at least several minutes before measuring the repassivation potential of the sample.

Weber and his coworkers [31] studied the aspects of dental corrosion on titanium system using various electrochemical techniques on titanium and its alloys with iron as an important constituent in dental media. The susceptibility to localized pitting corrosion of titanium and its alloys were evaluated by the breakdown potential, pitting potential, corrosion current density and the corresponding anodic polarization curves and tafel slopes.

Crevice corrosion

Localized crevice corrosion occurs from the geometry of the implant/prostheses assembly. Corrosion of an alloy is greater in the small sheltered volume of the crevice created by contact with another material. The other metal could be part of the fastener of the same or different alloy, a sheltered crown, cement packing or implant prostheses joint [32-33]. Crevice corrosion seems to prefer (metal-metal) Ti implant superstructure and base metal crown crevice with constricted space and oxygen gradient. Titanium alloy implants may be subjected to localized crevice attack exposed to short time periods of hot (> 70 [degrees]C, or 160 [degrees]F) chloride, bromide, iodide, fluoride or sulfate containing solutions during electrosurgery, electrocautery or thermocautery procedures. The reduction in pH and increase in crevicular chloride ion concentration are the essential factors in the initiation and propagation of the pits. When the acidity of the micro-environment around crevice increases with time it dissolves the passive oxide layer thereby causing localized destruction and crevice corrosion [34].

The mechanism for crevice corrosion in titanium is very similar to that seen in stainless steel, in which oxygen-depleted reducing acid conditions develop within tight crevices [32]. The model for crevice corrosion is illustrated in Figure 3. Dissolved oxygen or other oxidizing species in the bulk solution are depleted in the restricted volume of solution in the crevice. Finite surface oxidation in crevices consumes these species faster than diffusion from the bulk solution can replenish them as a result metal potentials within crevices become more active compared to metal surface exposed to the bulk body tissues. This situation creates a micro electrochemical cell in the bony socket where the shielded crevices become anode and corrode, and the surrounding more noble metal surface becomes cathode and is protected.

Titanium chlorides formed within the crevice are unstable and tend to hydrolyze, forming hydrochloric acid (HCl) and titanium oxide/hydroxide corrosion products. Because of the small, restricted volumes of solution within tight implant-superstructure crevice, low crevice pH levels (0-1) with high [H.sup.+] ions concentration can develop. These local reducing acidic conditions can result in severe and rapid localized active corrosion within crevices, depending on alloy resistance, temperature, and redox potential of the surrounding crevicular fluid.


Several oxidizing species such as oxygen, chlorine, ferric ion ([Fe.sup.3+]), and cupric ion ([Cu.sup.2+])from food and saliva, tend to effectively inhibit the general corrosion of exposed titanium implant surfaces, but most of these tend to aggravate the onset and propagation of titanium alloy crevice corrosion. These ionic species are excellent cathodic depolarizers and rate controlling, thus accelerate cathodic reduction kinetics involved in the electrochemical process. These cationic oxidizing species will not diffuse into the active crevice to inhibit attack compared to certain anionic oxidizing species such as N[O.sub.3.sup.-], Cl[O.sub.3.sup.-], O[Cl.sup.-], Cr[O.sub.3.sup.2-], Cl[O.sub.4.sup.-], and Mn[O.sub.4.sup.-], which can migrate into the implant-superstructure crevice inhibiting corrosion under halide solutions.

Crevice corrosion testing of titanium implants in function is insidious and very rapid, and may leach several ions into the crevicular space activating the host complement response and causing an adverse reaction that may or maynot be tolerated.

Several crevice test assemblies are employed including the multiple-crevice serrated washer and the sandwich-type crevice test assembly (consisting of 25 to 38 mm square flat sheet or plate specimen, with thin polytetrafloroethylene gaskets interspersed to provide the desired number of metal-to-metal and metal-to-gasket crevices) for the laboratory testing of crevice corrosion in titanium implants over varying time duration. In addition to visual and weight change measurements, monitoring of creviced implant potential and current has been used to a limited extent to identify the initiation of titanium alloy crevice corrosion. [35]

Environmentally induced cracking

Titanium alloy implants/prostheses processed under hydrogen-containing environments and under conditions in which galvanic couples or cathodic charging (impressed current) causes hydrogen to be evolved on metal surfaces may cause hydrogen induced cracking (HIC) [36-37] and the impairment of ductility of implant/prostheses permanently and irreversibly. The tenacious Ti[O.sub.2] surface film over the implant surface is a highly effective barrier to hydrogen penetration. Small amounts of moisture or oxygen in hydrogen-gas-containing environments effectively maintain this protective film, thus avoiding or limiting hydrogen uptake compared to anhydrous hydrogen atmospheres at higher temperatures and pressures causing increased uptake [38-39]. Hydrogen attack is the reaction of hydrogen with carbides in titanium to form methane, resulting in decarburization, surface voids and blisters. These voids are formed when the atomic hydrogen migrates from the surface to internal defects and inclusions, where molecular hydrogen gas can nucleate, generating internal stresses to deform and rupture the metal locally. In a and a-a alloys, excessive hydrogen uptake can induce the precipitation of titanium hydride in the a phase. These acicular-appearing hydride platelets, as shown in Figure 4 are brittle and have been well characterized in literature in detail [38].


Biaxial and triaxial mechanical properties such as ductility, cold-drawing/formability and impact toughness in a and near-a alloys are very sensitive to hydrogen levels [40-42]. Hydrogen contents above critical levels can result in sustained load cracking which drastically reduce the service life of cracked/notched or etched surfaces of implants under functional load of mastication and occlusion [40-42].Beta titanium alloys on the other hand have high solubility for hydrogen, therefore embrittlement is generally not associated with hydride precipitation. Severe loss in ductility and formability may not occur below several thousand parts per million of hydrogen [43]. This enhanced tolerance of beta alloys to hydrogen decreases with aging (tempering/ normalizing) as more of the a phase is precipitated. This increased a alloy tolerance must be weighed against higher hydrogen uptake rates that result from the higher hydrogen diffusion coefficient in a titanium alloys [43]. Covington et al. [39] have characterized the three conditions that must coexist to cause embrittlement of commercially pure titanium a implant alloys for dental applications in aqueous media.

1. A mechanism for generating nascent (atomic) hydrogen on a titanium surface. This hydrogen generation may be from a galvanic couple, an impressed cathodic current, corrosion of titanium, or severe continuous abrasion of the titanium surface in an aqueous medium.

2. Metal temperature above approximately 80 [degrees]C (175 [degrees]F), above which the diffusion rate of hydrogen into a titanium becomes significant [44]

3. Solution pH less than 3 or greater than 12, or impressed potentials more negative than -0.75 V (versus Ag/AgCl reference electrode) [45-47]

No hydrogen uptake and embrittlement problems occur when titanium implant is galvanically coupled to fully passive materials like titanium alloys, resistant (passive) stainless steels, copper alloys, and nickel-base alloys, depending on surrounding environments.

Very high cathodic charging of dental implant titanium alloys may cause enhanced hydride film formation and penetration over the implant surface and HIC at room temperature. Hot alkaline (> 80 [degrees]C, and pHe[]12), conditions may also result in increase hydrogen uptake and embrittlement of titanium alloys.

Hydrogen analysis for the titanium implants can be done using the electrochemical impedance spectroscopy and measuring the relative impedance of the retained hydrogen as a function of impedance or by the hot vacuum extraction method, where the implant samples are heated to 1100 to 1400 [degrees]C for a stipulated time to reversibly release the absorbed hydrogen, followed by evolved gas measurements.

Stress induced cracking

Stress-corrosion cracking (SCC) is a fracture phenomenon caused by the combined factors of tensile stress, a susceptible alloy, and a corrosive environment. The metal normally shows no evidence of general corrosion attack, although slight localized attack in the form of pitting may be visible in certain metals. Usually specific combinations of metallurgical and environmental conditions cause SCC. This combination of conditions is important because it is often possible to eliminate or reduce SCC susceptibility by modifying either the metallurgical characteristics of the metal and/or the environment. Another important aspect of SCC is the requirement for the tensile stress to be present, such as those developing from cold work, residual stresses during fabrication/machining, and externally applied functional/occlusal loads. Different surfaces of a metallic restoration (implant or crown structure) may have small pits and crevices and may be differentially exposed to different stresses consequently leading to stress corrosion cracking [36-37].The primary idea behind, titanium alloy SCC is the observation that no apparent corrosion, either uniform or localized, usually occurs before the cracking process [48-49] as a result it is difficult to translate the real oral situation into a laboratory experiment [49].

Over decades several models [49-54] have been proposed to explain the mechanism behind SCC in titanium alloys and broadly fall under two large categories.

1. Anodic-assisted cracking may begin where localized corrosion has occurred in the presence of a tensile stress. If corrosion is not rapid to impede the advancing crack tip, the crack will continue to advance into the metal and eventually lead to failure. Once a crack initiates, the balance among the crack tip corrosion rate, the crack tip environment, and the crack tip stress state is critical to crack propagation.

2. Hydrogen-assisted cracking (HIC) is said to occur by absorption of hydrogen near the crack tip. Hydrogen absorption leads to embrittlement of the metal ahead of the crack tip and promotes crack formation. The source of hydrogen is normally associated with anodic dissolution (that is, from the concurrent cathodic hydrogen-reduction reaction) at freshly exposed metal at the crack tip. As a result, anodic dissolution in the vicinity of the crack tip is normally required for this mechanism to operate.

ASTM has formulated the standard practice for testing SCC of titanium alloys in three basic categories [49]:

1. Use of smooth and statically loaded specimens such as U-bend, C-ring, bent beam, and dead-loaded tensile specimens.

2. Use of notched and pre-cracked specimens that are statically or dynamically loaded such as cantilever beam bend specimen, compact tension specimens and double-cantilever beam specimens.

3. Use of smooth or notched tensile specimens that are dynamically loaded at relatively low strain rates.

Fretting Corrosion/erosion corrosion

The combination of corrosive fluid (saliva with several enzymes and food particles) and high velocity in the oral environment results in erosion-corrosion or fretting. It is responsible for most of the metal release in tissue. Conjoint action of chemical (enzymes and proteins) and mechanical wear (mastication) during function further aggravates the attack [36-37,55-56]. In general during the passive environments, the hard and tenacious Ti[O.sub.2] surface film over the metal surface provides a superb barrier to erosion-corrosion. For this reason titanium alloys can withstand flowing water velocity as high as 30 m/s with little or no metal loss. The ability of the oxide film to repair itself when damaged and the intrinsic hardness of titanium alloys both contribute to their excellent resistance to erosion-corrosion.

Titanium alloys exhibit relatively high resistance to fluids containing suspended solids. Critical velocities for excessive metal removal depend upon the concentration, shape, size, hardness of the suspended particles, fluid impingement angle, [57] local turbulence, and titanium alloy properties.

The typically low concentrations of organic material in oral cavity is of little importance but continuous exposures to local changes around the implant during function can lead to finite removal of the metal as well as the cementing material between the implant and superstructure there by not only promoting erosion corrosion but crevice and galvanic corrosion as well.

Intergranular corrosion/cracking

For intergranular cracking to occur reactive impurities may segregate, or passivating elements may deplete at grain boundaries. This results in grain boundaries as a preferential region for corrosion owing to its high susceptibility and so the grains might fall out of the surface leading to material cracking [36-37]. This type of corrosion is not usually observed in titanium oral implants.

Galvanic corrosion of dental implants

The most common form of corrosion occurring in titanium implants is the galvanic corrosion or dissimilar corrosion. Titanium has been chosen as the material of choice for several trans-osseous and end-osseous implantations. Long term studies and clinical observations have established the fact that titanium is noble (due to the presence of adherent oxide) and does not corrode in human tissues however galvanic coupling of implant to several other metallic restorations may induce one of the several forms of corrosion. Thus coupling remains a great concern for the metallic superstructures covering the implant body.

Gold alloys have widely been used as a super structure owing to their excellent biocompatibility, corrosion resistance and mechanical properties. Owing to higher cost of the precious metal alloys (noble alloys)used in prosthodontics, it has led to the development of cost effective semi-precious metallic alternatives(non-noble alloys) [58-59] like gold-palladium, nickel-chromium, cobalt-chromium, nickel-titanium and several other titanium alloys.

Galvanic coupling occurs when two dissimilar metals are placed in direct contact within the oral cavity (adjacent or contralateral) or within the tissues. The complex corrosion process occurring involves the electrochemical reactions occurring at the dissimilar metals interface in presence of electrolyte (saliva or oral fluids or body fluids) resulting in the flow of electric current between them [37,60]. An in-vivo (in the living tissue) electrochemical cell is formed and galvanic current causes the corrosion of active metal and the noble metal is protected. The current also passes through the cellular junctions and tissues (desmosomes, hemi-desmosomes and cellular attachments) thereby activating the proprioceptors causing pain.

As already discussed in the section on hydrogen damage above attention should be given to possible excessive hydrogen uptake by titanium implant when it is galvanically coupled to active metal superstructures. This situation is of great concern in a titanium alloys when temperatures exceed 80[degrees]C in aqueous electrolytes during implant processing especially when hydrogen recombination poisons, such as sulfides, arsenides and cyanides are present.

Phenomena of galvanic corrosion

When two dissimilar metals with different electrode potentials come in contact in the presence of a corrosive electrolyte, a potential is generated. The net result is a chemical reaction with oxidation occurring at the anode and reduction occurring at the cathode, the electronic exchange occurs through the contact and ionic exchange occurs through the electrolyte.

The electrochemical cell reactions occurring at the different electrodes, depending upon the pH and aeration conditions as well the addition of oxidizers are [37]:

1) Anode (Oxidation)

M [R] [M.sup.n+] + [ne.sup.-]

2) Cathode (Reduction)

2[H.sup.+] + 2[e.sup.-][R] [H.sub.2]

4[H.sup.+] + 4[e.sup.-] + [O.sub.2] [R] 2[H.sub.2]O

2[H.sub.2]O + 2[e.sup.-][R] [H.sub.2] +2O[H.sup.-]

2[H.sub.2]O + 4[e.sup.-] + [O.sub.2] [R] 4O[H.sup.-]

[M.sup.n+] + [e.sup.-] [R] [M.sup.(n-1)+]

Thus if a base metal alloy superstructure is provided over the titanium implant, the less noble metal alloy forms the anode and the more noble titanium alloy is protected as a cathode. The electronic exchange occurs through the metallic contact and ionic contact occurs through saliva, mucosa, tissue and bone fluids. Though the hydrogen evolution at the implant surface may further complicate the situation by pressure under the prosthesis and hydrogen embrittlement of the implant surface.

Mechanisms of Corrosion resistance

The excellent corrosion resistance of titanium and its alloys used for implants is due the formation of a thermodynamically stable, continuous, highly adherent, and protective surface oxide film. Since titanium metal itself is highly reactive and has an extremely high affinity for oxygen, this beneficial surface oxide film is formed spontaneously and instantly when fresh metal surface is exposed to air and/or moisture. In fact, a damaged oxide film can generally reheal itself instantaneously if at least traces (that is, a few parts per million) of oxygen or water are present in the environment. Within a millisecond of exposure to air, a 10' oxide layer will be formed on the cut surface of the exposed pure Ti which will grow to about 100' thick within a minute.

The nature, composition, and thickness of the protective surface oxides that form on titanium alloys depend on environmental conditions. In most oral environments the oxide is typically Ti[O.sub.2] but may consist of mixtures of other titanium oxides as well including Ti[O.sub.2], [Ti.sub.2][O.sub.3], and TiO [61]. High-temperature oxidation promotes the formation of denser, more chemically resistant form of Ti[O.sub.2] known as rutile, whereas lower temperatures often generate a less crystalline and protective form of Ti[O.sub.2], anatase, or a mixture of rutile and anatase [61].

In dilute reducing acids, a 20 to 100 A multiplex film consisting of a hydrated titanium sesquioxide ([Ti.sub.2][O.sub.3]) inner layer and a Ti[O.sub.2] outer layer has been shown to form [62]. Increasing redox potential favors Ti[O.sub.2] formation, and increasing exposure temperature and/or time motivate conversion to the more stable crystalline form of Ti[O.sub.2]. These naturally formed films are typically less than 10 nm thick [63] and are not visible to the human eye. Ti[O.sub.2] oxide is highly resistant chemically and is attacked by very few chemicals like hot concentrated HCl, [H.sub.2]S[O.sub.4], NaOH, and HF. This thin surface oxide is also a highly effective barrier to hydrogen penetration of the alloy, as is discussed in a later section of this article. Furthermore, the Ti[O.sub.2] film being an n-type semiconductor exhibits increasing electronic conductivity with increasing temperature. As a cathode, titanium readily passes current and permits electrochemical reduction of ions in an aqueous electrolyte. On the other hand, very high resistance to anodic current flow (anodic polarization) across the passive oxide film can be seen in most aqueous solutions. Because the passivity of titanium occurs by the formation of a stable oxide film the corrosion behavior of implant metal interface can be understood by recognizing the conditions under which this oxide is thermodynamically stable.



The Pourbaix (potential-pH) diagram [64] for the titanium-water system at 25[degrees]C has been shown in Figure 5 and depicts the broad range around which the passive Ti[O.sub.2] film is stable, based on thermodynamic (free-energy) considerations. Oxide stability over the full pH scale is indicated over a wide range of highly oxidizing to mildly reducing potentials. Oxide film breaks and the corrosion of titanium implant surface occurs under reducing acidic conditions. Under strongly reducing (cathodic) conditions, titanium hydride formation is predicted. This range of oxide film stability and passivation is relatively insensitive to the presence of chlorides, explaining the high innate resistance of titanium to aqueous chloride environments [65].

The nature of the oxide film on titanium alloys remains unaltered in the presence of minor alloying constituents. Small additions (< 2 to 3%) of most commercially used alloying elements or variations in interstitial impurities have little effect on the corrosion resistance of titanium in normally passive environments. Increasing the alloying iron and sulfur content may increases corrosion rates exceeding ~0.10 mm/yr (3.9 mils/yr) [66]. Thus minor variations in alloy chemistry may be of concern only under conditions in which the passivity of titanium is borderline or when the metal is actively corroding.

Clinical significance of corrosion

Although titanium alloys have better corrosion properties compared to Co-Cr and stainless steel (other implant materials) their corrosion leads to dissolution of titanium and other alloying elements like aluminum, vanadium, niobium, molybdenum etc [67] causing localized to generalized host response as illustrated in Figure 6. The leached ions may induce potentially osteolytic cytokines into tissues leading to implant loosening [67] and may even cause severe allergic reactions or hypersensitivity [67]. Incidence of tumors and malignancies has also been reported in the literature but are few.

Fracture of Dental Implant

Fracture of dental implant/prostheses is a very rare phenomenon more often associated with mechanical function and previously use of screw preload systems to clamp flat to flat abutment implant junctions. They can have serious clinical complications. Corrosion can severely limit the fatigue strength and ultimate tensile strength of the material leading to its mechanical failure.

As reported by Green[68], the end-osseous implant superstructures, leached metal ions into the surrounding tissues as an event to corrosion, leading to fatigue fracture, following four years of functional loading into the oral cavity. Yokoyama et al. [69] on the other hand investigated the delayed fracture of titanium implant into the oral environment owing to hydrogen embritlement and environmentally induced cracking (EIC).

Bone loss and osteolysis

Olmedo, Fernandez and Guglidomotti [70] have speculated the corrosion induced ionic leaching to be responsible for peri-implantitis and treatment failure.

As already explained in Figure 6. the particles that are leached as a result of corrosion process are phagocytosized by macrophages (under the influence of host response) and release mediators of inflammation in the form of cytokines (host defense) [67] through the inflammatory cascade which inhibit the production of osteoblasts causing increased activity of the bone destroying cells leading to peripheral osteolysis and loosening of the implant.

Localized pain and inflammation (swelling)

Watterhehn et al. [71] have studied the effect of ionic release as a result of corrosion, causing localized pain and swelling with or without infection in the region of implant insertion. Figure 7 shows the localized inflammatory response to a freshly inserted implant in the right lateral incisor region (12 region). [71]

Cytotoxic tissue response

Corrosion release of the several substitutional alloying elements from various titanium alloys used in dentistry have been widely known in literature. Watterhehn et al. [71-76] have reported these metal ion release to be associated with carcinogenic and mutagenic activity of the oral cavity as shown in Figure 8.


Several studies have further shown that the cellular uptake of hexavalent chromium is many folds greater than the trivalent chromium ion and its increased uptake causes reduction of the alkaline phosphatase activity of the osteoblastic cells [77-79].

In-Vitro Studies

Several clinical cases of galvanic corrosion have been reported in the literature. Ravnholt et al. [80] studied the galvanic corrosion of titanium implants in contact with amalgam and cast metal alloys. As a continuation of their work Ravnholt and Jensen et al. [80-81] reported no changes in current or pH when titanium superstructures were in metallic contact with gold, cobalt-chromium, silver palladium and composite carbon alloys. Geis-Gerstorfer et al. [31] stated the importance of galvanic corrosion in implant system in two ways; first the possibility of biological effects like discoloration, pigmentation, localized reaction occurring from alloy components and secondly the osteolysis occurring as a mediation to phagocyte cascade activation.

Recalru and Meyer [60] studied the potential alloys that could be used as implant superstructure and concluded that an alloy that is potentially usable for superstructures in galvanic coupling to titanium must fill the following requisites

1. During coupling the titanium should have weak anodic polarization

2. The current generated during galvanism should be weak

3. The crevice potential must be much higher in comparison to common potential

Recently, Venugopalan and Lucas [82] defined the parameters for an acceptable couple combination to fulfill the following

1. The difference in the OCP (open circuit potential) and the corrosion current density (coupled materials) should be as small as possible

2. The breakdown potential of the anodic component should be significantly noble than the coupled corrosion density of the combination

3. The repassivation potential of the anodic component of the couple should be as similar as possible with minimal hysteresis


Johanson et al. [83] studied the effect of fluoride on corrosion behavior of machined and cast titanium with varying surface treatments and exposed area in contact with conventional and high copper amalgams. He observed higher current density for conventional amalgam in contact with titanium and concluded that surface preparations and anionic inclusions affect corrosion behavior of titanium implants.

In-Vivo Studies

Despite the use of titanium and its alloys as implants in human body increasing evidence is found that titanium and various substitutional alloying elements leach into the crevicular space around the implant [11-12,84-86].

Titanium like the other active-passive metals is covered by a protective air formed oxide of Ti[O.sub.2] which is adherent and stable. Although the oxide coating is thermodynamically stable, any form of chemical, electrochemical and mechanical trauma can lead to release of alloying elements through passive dissolution process.

Ferguson and coworkers [87] first documented the elevated titanium levels locally around the implant periphery. Bumgardner et al. [88] reported the increased release of metal ions through a gallium titanium galvanic couple causing increased implant and tissue deterioration.Cortada et al. [89] used the ICP-AES (inductively coupled plasma-atomic emission spectroscopy) technique to confirm the release of metallic ions during titanium implant galvanism.

Nakagawa et al. [90] studied the effect of fluoridation during prophylactic therapy on the corrosion behavior of an implant and reported higher corrosion current density with increasing anionic concentrations. Extended work has been done by Bhola et al. [91] to study the effect of Knutson fluoridation on the corrosion behavior of titanium-niobium dental implant alloys in normal saline.

Kirkpatrick et al. [92] highlighted the patho-physiology and patho-mechanics behind impaired wound and tissue healing under metallic ions release during corrosion.

Dental corrosion Mitigation

With the development of corrosion science and advances in material engineering and design several strategies have been employed today to prevent in-use (within the oral cavity) and out-use (outside oral cavity) corrosion of titanium implants and prostheses.

1. Alloying titanium with noble metals (platinum-group metals) which facilitates cathodic depolarization by catalyzing the hydrogen reduction step in acid solutions used during active processing of dental implants in the industry for out-use prevention.

2. Alloying titanium with more thermodynamically stable acid-resistant elements such as molybdenum, zirconium, tantalum, niobium used during processing of dental implants, also makes the implant surface more biocompatible and cell adherent for in use prevention.

3. Addition or presence of soluble oxidizing ions or compounds (cathodic depolarizers) in the oral media (e.g., [Fe.sup.3+], [O.sub.2], [Cl.sub.2], N[O.sub.3.sup.-]), can be used during implant insertion and placement using NS, RL, PBS etc, for in-use prevention.

4. Noble metal surface treatments (platinum-group metal coatings or enriched/modified surfaces) can be used during implant surface processing for in-use corrosion prevention.

5. Anodic protection via impressed positive potentials from an external power source or from direct galvanic coupling with a noble metal (platinum-group metal) can be used during processing for out-use prevention.

6. Thermal oxidation or nitriding of titanium surfaces can be used during implant manufacturing and processing to facilitate in-use prevention.

To summarize several techniques do exist to effectively reduce the corrosion in dental implants, but we can mainly rely on material selection and coating the surface for an effective in-use clinical performance into a patient's oral cavity.


The metallic titanium dental implants/prostheses used in dentistry today derive their biocompatibility from the alloying elements responsible for the formation of a continuous stable Ti[O.sub.2] passive film on its surface.

There is a significantly small release of alloying ions even under the ideal conditions of passivity and with no damage to the implant surface. Corrosion of these implants may occur when the oral conditions are unfavorable as under mechanical trauma to the implant surface (during placement, subject induced, and trauma to assault) or the use of inappropriate metal combination as auxiliary prostheses (galvanism).

The potential adverse effects of metal ion release into living tissues can be proposed based on information from literature and various clinical, preclinical and animal trial studies in-vivo and in-vitro. The results of in-vivo and in-vitro testing do not necessarily take into account all of the protection mechanisms and physiological host response characteristics of the actual implant environment in oral cavity.

It is clear that corrosion is bound to occur and its consequences can be quite severe. Our current research standards, regulations for biocompatibility testing, material and design understanding and composition of metals and its alloys suitable to a special application have drastically reduced the incidence of implant failures in oral physiological environment under corrosion and its adverse events.


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Rahul Bhola *, Shaily M. Bhola, Brajendra Mishra and David L. Olson

Dept of Metallurgical & Materials Engineering, 1500 Illinois Street, Golden, Colorado 80401, USA

* Corresponding author ( Bhola

Received 31 July 2010; Accepted 3 August 2010; Available online 1 March 2011
Table 1: Metallic biomaterials for surgical implant and
prostheses. [4-6]

Metallic Biomaterials for Implants/Prostheses

Material CommonName/
Designation Trade Name

Speciality Metallic Biomaterials

Ta, Unalloyed, cast Unalloyed Tantalum ([alpha])
Zr-2.5 Nb Zircadyne
Ni-45Ti Nitinol

Titanium Base Biomaterials

Ti CP-1 CP-1 ([alpha])
Ti CP-2 CP-2 ([alpha])
Ti CP-3 CP-3 ([alpha])
Ti CP-4 CP-4 ([alpha])
Ti-3Al-2.5V Ti-3Al-2.5V ([alpha]/[beta])
Ti-5Al-2.5Fe Tikrutan ([alpha]/[beta])
Ti-6Al-4V Ti-6Al-4V ([alpha]/[beta])
Ti-6Al-4V cast Ti-6Al-4V ([alpha]/[beta])
Ti-6AUV ELI Ti-6Al-4V ELI ([alpha]/[beta])
Ti-6Al-7Nb Ti-6Al-7Nb ([alpha]/[beta])
Ti-15Mo Ti-15Mo (Metastable [beta])
Ti-12Mo-6Zr-2Fe TMZF (Metastable [beta])
Ti-11.5Mo-6Zr-4.5Sn Beta 3 (Metastable [beta])
i-13Nb-13Zr Ti-13Nb-13Zr (Metastable [beta])
Ti-45Nb Ti-45Nb (Metastable [beta])
Ti-35Nb-7Zr-5Ta TiOsteum (Metastable [beta])

Cobalt Base Biomaterials

Co-28Cr-6Mo casting alloy Cast CoCrMo
Co-28Cr-6Mo wrought alloy #1 Alloy 1
Co-28Cr-6Mo wrought alloy #2 Alloy 2
Co-28Cr-6Mo wrought alloy #3 GADS
Co-20Cr-15W-10Ni-1.5Mn L-605
Co-20Cr-20Ni-5 Fe-3.5Mo-3.5W-2Ti Syncoben
Co-19Cr-17Ni-14Fe-7Mo-1.5Mn Grade 2-Phynox
Co-20Cr-15Ni-1 5Fe-7Mo-2Mn Grade 1-Elgiloy
Co-35Ni-20Cr-1 0Mo 35N

Steel Base Biomaterials

Fe-18Cr-14Ni-2.5Mo 316 L Stainless Steel
Fe-18Cr-12.5Ni-2.5Mo cast 31 6 L Stainless Steel
Fe-21Cr-10Ni-3.5Mn-2.5Mo REX 734
Fe-22Cr-12.5Ni-5Mn-2.5Mo XM 19
Fe-23Mn-21Cr-1Mo-1N 108

Material UNS Design- ASTM
Designation ation Designation

Speciality Metallic Biomaterials

Ta, Unalloyed, cast R05200 F 560
Zr-2.5 Nb R60705 F 04.12.45
Ni-45Ti N01555 F 2063

Titanium Base Biomaterials

Ti CP-1 R50250 F 67
Ti CP-2 R50400 F 67
Ti CP-3 R50550 F 67
Ti CP-4 R50700 F 67
Ti-3Al-2.5V R56320 F 2146
Ti-5Al-2.5Fe Unassigned -
Ti-6Al-4V R56400 F 1472
Ti-6Al-4V cast R56406 F 1108
Ti-6AUV ELI R56401 F 136
Ti-6Al-7Nb R56700 F 1295
Ti-15Mo R58150 F 2066
Ti-12Mo-6Zr-2Fe R58120 F 1813
Ti-11.5Mo-6Zr-4.5Sn R58030 F 9046
i-13Nb-13Zr R58130 F 1713
Ti-45Nb R58450 F 04.12.44
Ti-35Nb-7Zr-5Ta R58350 F 04.12.23

Cobalt Base Biomaterials

Co-28Cr-6Mo casting alloy R30075 F 75
Co-28Cr-6Mo wrought alloy #1 R31537 F1537
Co-28Cr-6Mo wrought alloy #2 R31538 F 1537
Co-28Cr-6Mo wrought alloy #3 R31539 F 1537
Co-20Cr-15W-10Ni-1.5Mn R30605 F 90
Co-20Cr-20Ni-5 Fe-3.5Mo-3.5W-2Ti R30563 F 563
Co-19Cr-17Ni-14Fe-7Mo-1.5Mn R30008 F 1058
Co-20Cr-15Ni-1 5Fe-7Mo-2Mn R30003 F 1058
Co-35Ni-20Cr-1 0Mo R30035 F 562

Steel Base Biomaterials

Fe-18Cr-14Ni-2.5Mo S31673 F 138
Fe-18Cr-12.5Ni-2.5Mo cast Unassigned F745
Fe-21Cr-10Ni-3.5Mn-2.5Mo S31675 F 1586
Fe-22Cr-12.5Ni-5Mn-2.5Mo S20910 F 1314
Fe-23Mn-21Cr-1Mo-1N Unassigned F 04.12.35

Material ISO Alloy #
Designation Designation

Speciality Metallic Biomaterials

Ta, Unalloyed, cast - 0.0
Zr-2.5 Nb - 2.5
Ni-45Ti - 55.0

Titanium Base Biomaterials

Ti CP-1 ISO 5832-2 0.1
Ti CP-2 ISO 5832-2 0.2
Ti CP-3 ISO 5832-2 0.3
Ti CP-4 ISO 5832-2 0.4
Ti-3Al-2.5V - 5.5
Ti-5Al-2.5Fe ISO 5832-10 7.5
Ti-6Al-4V ISO 5832-3 10.0
Ti-6Al-4V cast - 10.1
Ti-6AUV ELI ISO 5832-3 10.2
Ti-6Al-7Nb ISO 5832-11 13.0
Ti-15Mo - 15.0
Ti-12Mo-6Zr-2Fe - 20.0
Ti-11.5Mo-6Zr-4.5Sn - 22.0
i-13Nb-13Zr - 26.0
Ti-45Nb - 45.0
Ti-35Nb-7Zr-5Ta - 47.0

Cobalt Base Biomaterials

Co-28Cr-6Mo casting alloy ISO 5832-4 34.0
Co-28Cr-6Mo wrought alloy #1 ISO 5832- 12 34.1
Co-28Cr-6Mo wrought alloy #2 ISO 5832-12 34.2
Co-28Cr-6Mo wrought alloy #3 - 34.3
Co-20Cr-15W-10Ni-1.5Mn ISO 5832-5 46.5
Co-20Cr-20Ni-5 Fe-3.5Mo-3.5W-2Ti ISO 5832-8 49.0
Co-19Cr-17Ni-14Fe-7Mo-1.5Mn ISO 5832-7 58.5
Co-20Cr-15Ni-1 5Fe-7Mo-2Mn ISO 5832-7 59.0
Co-35Ni-20Cr-1 0Mo ISO 5832-6 65.0

Steel Base Biomaterials

Fe-18Cr-14Ni-2.5Mo ISO 5832-1 34.5
Fe-18Cr-12.5Ni-2.5Mo cast - 34.6
Fe-21Cr-10Ni-3.5Mn-2.5Mo ISO 5832-9 37.0
Fe-22Cr-12.5Ni-5Mn-2.5Mo - 42.0
Fe-23Mn-21Cr-1Mo-1N - 46.0

Figure 1: Prerequisites for an implant biomaterial [17]



* Tissue * Fabrication
 reactions * Elasticity methods.

* Changes in * Yield stress * Consistency
 properties and conformity
 * Ductility to all
 * Mechanical requirements
 * Toughness
 * Physical * Quality of
 * Time raw materials
 * Chemical dependent
 deformation * Superior
 techniques to
* Degradation * Creep obtain
 leads to excellent
 * Ultimate surface finish
 * Local strength or texture
 deleterious * Capability
 changes * fatigue of material to
 strength get safe and
 * Harmful efficient
 systemic * Hardness sterilization
 * Wear * Cost of product
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Author:Bhola, Rahul; Bhola, Shaily M.; Mishra, Brajendra; Olson, David L.
Publication:Trends in Biomaterials and Artificial Organs
Date:Jan 1, 2011
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