Electrical Implications of Corrosion for Osseointegration of Titanium Implants

Abstract

The success rate of titanium implants for dental and orthopedic applications depends on the ability of surrounding bone tissue to integrate with the surface of the device, and it remains far from ideal in patients with bone compromised by physiological factors. The electrical properties and electrical stimulation of bone have been shown to control its growth and healing and can enhance osseointegration. Bone cells are also sensitive to the chemical products generated during corrosion events, but less is known about how the electrical signals associated with corrosion might affect osseointegration. The metallic nature of the materials used for implant applications and the corrosive environments found in the human body, in combination with the continuous and cyclic loads to which these implants are exposed, may lead to corrosion and its corresponding electrochemical products. The abnormal electrical currents produced during corrosion can convert any metallic implant into an electrode, and the negative impact on the surrounding tissue due to these extreme signals could be an additional cause of poor performance and rejection of implants. Here, we review basic aspects of the electrical properties and electrical stimulation of bone, as well as fundamental concepts of aqueous corrosion and its electrical and clinical implications.

Biopotentials

Exogenous electrical control of cell and tissue physiology has been studied since the late 1700s with the work of scientists such as Luigi Galvani, Alessandro Volta, Carlo Matteucci, and Emil Du-Bois Reymond, and the discovery of biopotentials and injury potentials (Black, 1986; Piccolino, 1997). Biopotentials are natural electrical properties that control normal growth and development of different types of cells and tissues (Ferrier et al., 1986; Levin et al., 2002) (Figs. 1a, 1b). Injury potentials are alterations to the normal potential patterns of intact tissue (Becker et al., 1977; Levin, 2007), characterized by stable, long-lasting direct current (DC) voltage potentials induced between injured and intact tissues that persist until the wound has healed. These potentials can span hundreds of microns and are generated by current or ions flowing through the injured tissue (Lokietek et al., 1974; McCaig et al., 2005) (Figs. 1c, 1d). Currents of 1-100 µA/cm² have been measured in injured tissues (Lokietek et al., 1974; Borgens et al., 1980), and, assuming the resistivity of soft tissues to be 100 Ω·cm (Faes et al., 1999; McCaig et al., 2005), these currents create voltage differences of 10-100 mV/cm across hundreds of microns.

Figure 1.

Schematic shows: (a) electrical potential of a cell across an intact plasma membrane (V_m), and (b) the inward current flow, and its respective potential (V) after localized injury to the plasma membrane; (c) transepithelial electrical potential (V_TEP) across an intact cell layer of the skin; and (d) short-circuit caused by a wound. Adapted from McCaig et al. (2005).

Recent findings underscore the importance of endogenous electrical potentials in cell signaling and gene expression. Endogenous electrical potentials, and specifically injury potentials, have been associated with epithelial cell migration and advancement of the wound-healing front through activation of Src and inositol-phospholipid signaling pathways in a rat corneal model (Zhao et al., 2006). Disruption of endogenous electric potentials affected the migration speed and direction of the wound-healing front. The same group also found that corneal epithelial cells from bovine eyes were sensitive to directional cues such as nanogrooves (i.e., contact guidance) and electric fields (i.e., electrotaxis) through the activation of small GTPases, rho and cdc42, respectively (Rajnicek et al., 2007). The study showed that electrotaxis seemed to be more potent than, but not completely dominant over, contact guidance by setting the electric fields orthogonally to the nanogrooves and measuring the distance traveled by the cells. Furthermore, a cell-membrane voltage sensor, Ciona intestinalis voltage-sensor-containing phosphatase (Ci-VSP), has been identified, which is activated by changes in membrane potential and can initiate signaling cascades (Murata et al., 2005; Iwasaki et al., 2008).

Electrical Signals in Bone

Both biopotentials and injury potentials are found in bone. Electrical properties and electrical stimulation of bone have been investigated since the 1950s, beginning with the piezoelectric nature of osseous tissue (Fukada and Yasuda, 1957). When forces were applied to sections of previously dried human and ox femurs, directly proportional voltages could be measured that were dependent on the collagen fiber alignment. This led to the idea that electrical signals could be related to the process of bone formation. Additional endogenous electrical properties of bone have been discovered since and are suggested to play a role in the feedback mechanism of bone remodeling and development (Guzelsu and Demiray, 1979; Rubinacci et al., 1988).

Biopotentials in bone are classified into two subgroups, due in part to the complexity of bone structure: strain-related potentials (SRP) and biopotentials. SRPs include the piezoelectric behavior (i.e., electric potential in response to applied forces) of bone due to the structure and dipolar charge of collagen, and streaming potentials associated with the flow of fluid and ions through porous bone. The subgroup of biopotentials in bone results from contribution of biological processes such as osteoblast membrane potential, extracellular matrix acidification and ion release caused by osteoclast bone resorption, and cell junctions of osteocytes. In vivo, these electrical signals work in concert to provide the correct environment for normal bone growth and development, but can be disrupted or altered by injury potentials after trauma and during healing.

Mechanical forces have been shown to direct the process of bone remodeling (Burr et al., 2002; Hou et al., 2007). Accordingly, areas of bone under stress tend to grow, and those areas under no mechanical load tend to be resorbed (Duncan and Turner, 1995). This is believed to be a result of the physical stress alteration and biochemical activation of particular bone cells (Duncan and Turner, 1995). As a parallel event, however, areas of bone that are under mechanical load generate a more negative polarity than areas under smaller or no loads (Fukada and Yasuda, 1957; Bassett and Becker, 1962) (Fig. 2a). Thus, bone growth could also be attributed to negative polarity and bone resorption to positive polarity, suggesting that electrical signals work as a feedback mechanism for bone remodeling (Becker et al., 1977; Black, 1986).

Figure 2.

Schematic of: (a) polarization of bone under applied mechanical forces; (b) voltage vs. distance comparison between intact and one-hour post-fracture bones; and (c) simplified mechanical forces and respective polarization of bone and periodontal ligament under orthodontic treatment. Adapted from Black (1986).

The relationship between negative potentials and bone growth is seen during long-bone fracture-healing (Bassett et al., 1964; Becker et al., 1977; Black et al., 1984). In children, fractured long bones tend to overgrow with respect to their counterparts (Aitken et al., 1939; Wray and Goodman, 1961), and there is an increase in apoptosis in the growth plate (Gaber et al., 2009). Interestingly, both the healing site and growth plate tend to have a more negative potential compared with that of the nearby intact tissue (Friedenberg and Brighton, 1966; Black, 1986) (Fig. 2b). During development, the growth plate has a negative potential, while the growth plate of a mature individual tends to have a neutral voltage (Lai et al., 2000). Consequently, negative potentials in the growth plate after fracture may be related to bone overgrowth, since cortical bone healing and repair should not increase bone length. However, the negative potential found at the fracture site may coordinate bone healing and may directly influence the polarity of the growth plate.

The relationship among mechanical stimuli, electrical potentials, and bone remodeling can also be seen in orthodontic treatment of individuals with malocclusion. Tipping and translational mechanical forces applied during orthodontic treatment deform and remodel alveolar bone and periodontal tissue, resulting in tooth movement (Isaacson et al., 1993; Meikle, 2006). Quantitative techniques such as finite element analysis have been used to assess forces affecting tooth movement (McGuinness et al., 1991; Cobo et al., 1993). Although the experimental design and parameters measured varied between and among studies, the results indicate a direct relationship between the magnitude of the applied stress and the level of bone and periodontal ligament remodeling. Some studies have correlated excessive forces to orthodontically induced inflammatory root resorption (Brezniak and Wasserstein, 2002).

Despite innate differences between the origin of long bones and maxillofacial bone, forces applied on teeth and surrounding alveolar bone generate similar electrical potentials (Cochran et al., 1967; Zengo et al., 1973). The electrical polarization of these tissues has also been correlated to bone remodeling. Areas with high osteoblast activity and bone growth show negative polarization, while areas under resorption due to higher osteoclast levels show a positive or neutral polarization (Norton et al., 1984) (Fig. 2c). One suggested hypothesis states that these electrical potentials may provide a more direct measurement of the mechanical forces delivered by orthodontic devices and help provide a more personalized treatment (Norton, 1975).

Electrical Stimulation of Bone

The roles of these electrical signals in bone growth and development have prompted several research groups to study bone repair using methods to stimulate cells and tissues electrically in vitro (Wang et al., 1998; Kim et al., 2006) and in vivo (Brighton and Hunt, 1986; Shafer et al., 1995) with very successful outcomes. Some have associated bone growth enhancement after electrical stimulation with the production of osteoinductive factors such as bone morphogenetic proteins (BMPs) 2, 4, 5, and 6 (Kim et al., 2006; Wang et al., 2006), as well as with levels of intracellular and extracellular calcium (Wang et al., 1998; Aaron et al., 2004). However, differences in experimental design and the electrical parameters used, and the over-simplification of in vitro models that do not account for many aspects of in vivo conditions, have hindered the systematic investigation of the molecular pathways involved with cell responses. Additionally, it is not well-understood if electrical stimulation affects every type of bone (e.g., cortical, trabecular, membranous) similarly. Fundamental understanding of the molecular pathways involved in electrical stimulation is necessary to elucidate the role of electrical signals in the bone-implant interface, thus allowing for better system designs for personalized treatment.

Electrical stimulation systems can be classified into 3 main groups, depending on the nature of the electrical signals being supplied: direct current stimulation, capacitive stimulation, and inductive stimulation.

Direct Current (DC) Stimulation

DC stimulation, or faradic stimulation, is an invasive method that applies a DC electric field to growing cells, either directly through the surface on which they are growing, or indirectly through the medium in which they are growing (Fig. 3a). Common parameters applied include fixed currents of 1-50 µA/cm², which can affect osteoblast proliferation and expression of differentiation markers (Bodamyali et al., 1999; Kim et al., 2006). The majority of published studies of in vitro DC stimulation used electrodes submerged in the tissue culture medium, establishing a DC electric field and inducing electrochemical currents between the anode and the cathode (Curtze et al., 2004; Ercan and Webster, 2008). However, the products generated at the cathode and the anode that have enhancing or detrimental effects on cell response, respectively (Black et al., 1984; Bodamyali et al., 1999), may obscure the results of DC electrical stimulation.

Figure 3.

Schematics of different electrical stimulation systems. (a) DC electrical stimulation set-up consisting of a battery that generates an electric field (EF) directly through the implant device. The implant becomes the cathode, the anode is exposed to the oral cavity, and the surrounding tissue serves as a path to close the circuit and allow for the flow of current. (b) Capacitive stimulation set-up, consisting of 2 externally applied electrodes that generate an electric field (EF). (c) Inductive stimulation set-up, consisting of a pair of multi-turn Helmholtz coils connected in series to generate an electromagnetic field (EMF).

Titanium (Ti) implants can be used as cathodes for DC electrical stimulation (Dodge et al., 2007). One such device was developed to fit inside a dental implant healing abutment and supply electrical stimulation to canine mandibular bone (Song et al., 2009). Biphasic electrical stimulation increased bone formation and bone-to-implant contact when compared with control implants. Although Ti is one of the most-used materials for bone implants, to our knowledge there are no in vitro studies evaluating osteoblast response to electrical stimulation when grown directly on a Ti cathode. Ti substrates used in a typical configuration with submerged electrodes in the media showed increased osteoblast density by DAPI staining (Ercan and Webster, 2008, 2010).

Capacitive Stimulation

In capacitive stimulation, electrodes are applied externally to the skin above the area to be stimulated, inducing an electric field that can influence cell response (Fig. 3b). Common external stimulation systems use alternating current (AC) parameters that vary between 1 and 50 V at frequencies of 60-200,000 Hz, and effective electric field strengths from 0.1 to 5 V/m (Lorich et al., 1998; Shigino et al., 2001). Cells grown in vitro on tissue culture substrates are either stimulated through the media with an AC power supply or sandwiched between electrodes without media contact. Capacitive stimulation is advantageous because it is non-invasive, and it has been shown to have an effect both in vitro (Brighton et al., 2001; Wang et al., 2006) and in vivo (Brighton et al., 1985; Lirani-Galvao et al., 2009). However, the therapeutic results depend on patient compliance, and high voltages and frequencies applied may cause irritation (Black, 1986; Gan and Glazer, 2006). Since capacitive stimulation cannot be applied directly to the affected osseous tissue, and because of the complexities of measuring local current densities in the site of interest, it is difficult to predict the subsequent effects.

Inductive Stimulation

Inductive stimulation is a non-invasive method that uses a coil or pair of coils connected in series, with their axes perpendicular to the long bone, to generate pulsed electromagnetic fields (PEMFs) and small secondary electric fields (Black, 1986) (Fig. 3c). These magnetic fields and the induced electrical fields have been shown to influence cell response and gene expression (Bodamyali et al., 1998; Brighton et al., 2001; Schwartz et al., 2008).

In a series of studies performed in our laboratory, the effects of PEMFs on MG63 osteoblast-like cells were shown to reduce cell number, and increase osteoblast maturation, collagen synthesis, and local factor production, including transforming growth factor-β1 (TGF-β1) (Lohmann et al., 2000b). Cells from human hypertrophic and atrophic non-union tissues have been used to evaluate the effects of PEMFs on non-union fractures, commonly treated with electrical stimulation (Guerkov et al., 2001). Cells exposed to PEMFs increased TGF-β1, with no effect on cell proliferation or differentiation, suggesting that improvements in non-unions after PEMFs result from changes in local factor production near the affected area. Finally, mesenchymal stem cells (MSCs) have been used to evaluate the effects of PEMFs on progenitor cell differentiation, one of the first types of cells to arrive after implant placement (Schwartz et al., 2008, 2009). PEMF synergistically increased MSC osteogenesis when cells were cultured on calcium phosphate disks in the presence of the osteoinductive factor bone morphogenetic protein-2 (BMP-2), as determined by increased alkaline phosphatase, osteocalcin, and TGF-β1. These results suggest that electrical stimulation may also improve bone healing and osseointegration by increasing osteogenic differentiation of MSCs.

Like capacitive stimulation, inductive stimulation has no electrochemical effect on the tissue, because it is non-invasive. Clinically, one disadvantage is that therapy success depends on patient compliance (Gan and Glazer, 2006). Additionally, the non-localized application of inductive stimulation may affect multiple types of tissues surrounding the injury site.

Taking into consideration what is known about mechanical stimulation and the different electrical stimulation systems, improvements in bone growth and repair can be achieved through different pathways, including integrin- and IP3-mediated pathways (Wang et al., 1993; Brighton et al., 1996) (Fig. 4).

Figure 4.

Schematic of possible signaling pathways used by mechanical and electrical stimulation systems.

Electrical Implications of Corrosion

Metals are used for dental and orthopedic implants because of their mechanical properties, such as weight-to-strength ratio and good biological performance. However, metallic devices are prone to corrosion, particularly in aqueous environments under extreme conditions. Corrosion resistance depends on temperature, pH, ion concentration, substrate size, and chemistry, but it is not inherent to the material itself as has been implied in many studies (Bhattarai et al., 2008). Ti is corrosion-resistant under controlled environments in the absence of load. In the human body, the physiological environment in combination with constant, cyclic implant loading can significantly enhance corrosion rates (Long and Rack, 1998; Papakyriacou et al., 2000; Brunette et al., 2001; Lewis et al., 2005). Extreme acidic conditions found during inflammation (Lassus et al., 1998), fretting between implant and bone (Gilbert et al., 2009), and galvanic corrosion between Ti implants and other metallic alloys used for common dental procedures (Grosgogeat et al., 1999) could greatly affect the mechanical stability and outcome of dental implants.

Basic Electrochemistry

The basic unit of electrochemistry is the electrochemical cell, which is composed of 2 electrodes (anode and cathode) and an aqueous electrolyte serving as a connecting path. Electrochemical reactions on the surface of an electrode can be oxidative (anodic), generating electrons and ions, or reductive (cathodic), consuming electrons and generating metal atoms or other molecules (Fig. 5). An electrode is defined by how reactive it is compared with the opposite electrode to which it is connected. In some situations, a single metal device can serve as both the anode and the cathode, and so a second electrode is not required to complete the circuit.

Figure 5.

Schematic of initiation and mechanism of corrosion of a dental implant.

Metallic implants for bone applications submerged in ion-rich electrolytes in the body constitute a basic corrosion cell. Large currents can be induced by the flow of ions and electrons generated during electrochemical reactions occurring between the corroding metallic surface and the electrolyte. These currents are generally used to measure the corrosion rate of a metal, because they are directly related to the release of metal ions or, in other words, to the material’s degradation. Consequently, corrosion events result in the formation of small pits on the surface of the device that can amplify the corrosive environment around the implant and compromise its mechanical stability. This can lead to shortening of the implant’s lifetime and sudden failure (Papakyriacou et al., 2000; Teoh, 2000; Mudali et al., 2003) (Fig. 5). Products of the electrochemical reactions may have cytotoxic or even neoplastic effects on the tissue surrounding the implant, serving as an additional cause of rejection or aseptic loosening (Doran et al., 1998; Denaro et al., 2008b; McGuff et al., 2008). However, the electrical implications of corrosion on the surrounding tissue have not been extensively investigated (Denaro et al., 2008a).

Passivity of Titanium

Certain metals like Ti oxidize easily, forming a very thin, stable passive layer that is self-limiting and protects the surface of the metal from further oxidation. This behavior, passivity, gives Ti its high corrosion resistance under certain controlled conditions where, otherwise, it would undergo strong active corrosion. Metals can have stable passivity, where the oxide layer self-heals immediately after being ruptured, or unstable passivity, where the oxide layer is unable to heal after disruption and the bare metal is exposed to active corrosion. Both of these events depend on the oxidizing or reducing potential of the environment. The passive oxide layer formed on the surface of Ti may be responsible for its good biological performance, since it is less reactive than bare Ti. Additionally, it may mimic the ceramic nature of bone and allow biochemical bonding with the newly formed bone (Sul et al., 2005).

Most materials chosen for implant applications exhibit passivity properties and, thus, relatively low corrosion rates compared with those of other more reactive metals, such as zinc, magnesium, or vanadium, which undergo active corrosion even in relatively neutral pH. However, certain environmental conditions can breach the protective oxide layer formed on the surfaces of these passive materials and cause corrosion, affecting the mechanical integrity of the implant and the health of the surrounding tissue.

Work by our laboratory and many others has shown that implant surface properties such as roughness, chemistry, and energy directly influence tissue response by affecting protein adsorption and modulating cell proliferation and differentiation (Schwartz and Boyan, 1994; Kieswetter et al., 1996). Additionally, innovations in surface modification techniques have improved the biological performance of metallic implants (Wang et al., 2003; Buser et al., 2004). However, some modifications may diminish mechanical properties of the bulk material, resulting in surface micro-cracks, increased corrosion rates (Papakyriacou et al., 2000; Teoh, 2000; Hazar Yoruc and Kelesoglu, 2009), and, thus, increased corrosion currents and potentials that may affect surrounding cells and tissues.

Types of Corrosion

The most common types of corrosion found in metallic materials used for implant applications are galvanic, fretting, and pitting/crevice corrosion, as well as environmentally induced cracking (EIC).

Galvanic corrosion occurs with direct contact of two dissimilar metals in an electrolytic solution (Kaesche, 2003). The difference in electrochemical potential of the two metals promotes oxidation of the more reactive metal. This becomes the anode, which generates a flow of electrons and ions to the cathode. In one study, spine implant constructs consisting of pedicle screws, connectors, and rods that had mixed components made of stainless steel (SS) and titanium were investigated for signs of galvanic corrosion under dynamic loads (Serhan et al., 2004). The results showed no evidence of corrosion on surfaces of the implant that had not been in contact with other components, and only minor signs of corrosion at the interfaces between SS-Ti, Ti-Ti, and SS-SS, with the latter actually having the greatest amount of corrosion. Galvanic corrosion is not common in dental implant applications because of the presence of only one component, the dental screw, and the insulating nature of the protective passive layer that forms on the surface. Nevertheless, in some individuals the surrounding tissue could serve as a medium for electrical flow between metallic implants and other types of alloys used in dentistry for amalgams or orthodontic devices. Galvanic corrosion could also amplify the rates of corrosion initiated by other mechanisms described below (Reclaru and Meyer, 1994; Grosgogeat et al., 1999).

Fretting corrosion is caused by the repeated micro-motion or friction of a metal component against another material that causes mechanical wear and breaks up the passivating layer on the contact surface of the metallic device (Landolt, 2007). Fretting between dental implants and bone during implantation and due to cyclic loads imparted from chewing has been suggested as a cause of Ti corrosion and metal ion release (Denaro et al., 2008b; Gilbert et al., 2009). Fretting could also be an issue in total hip replacements, where it could generate wear debris and ions from friction between joint and socket (Long and Rack, 1998; Ingham and Fisher, 2000). The release of metal debris and ions has been linked to inhibition of cell differentiation, cytotoxicity, phagocytosis of Ti particles by macrophages and other cells, inflammation, and neoplasia (Sun et al., 1997; Doran et al., 1998; Lohmann et al., 2000a; Rahal et al., 2000). Recent studies have shown that fretting and oxide disruption at the surface of load-bearing implants can cause corrosion current densities to increase and generate open-circuit potentials in excess of -500 mV (Goldberg and Gilbert, 1997; Gilbert et al., 2009). Abnormal electrical signals may affect the response and stability of the adjacent tissue, and fretting corrosion may amplify other types of corrosion by rupturing the passivating film and exposing bare Ti.

Pitting corrosion occurs as a result of the spontaneous breakdown of the passive film on a flat and evenly exposed area (Evans, 1960; Landolt, 2007). Crevice corrosion is a localized corrosion due to a geometric confinement in the design of the device or from a previously corroded region on the surface. The common mechanism of propagation for both usually involves a differential aeration cell (Fig. 5). In this, the region undergoing active corrosion has restricted solution flow due to geometric confinement and initially depletes local oxygen concentration, generating high levels of metal ions and electrons that are consequently consumed by the surface exposed to high levels of oxygen. While pitting corrosion is not likely to occur on Ti surfaces, crevice corrosion has been found (Charles and Ness, 2006; Denaro et al., 2008b). In one study, corrosion currents from Ti alloy lumbar interbody fusions were directly related to lumbar pain and periprosthetic bone loss in patients (Denaro et al., 2008b).

EIC is the brittle mechanical failure of metallic devices under stress levels significantly lower than their ultimate tensile strength. This occurs in susceptible materials in corrosive environments and under continuous loading. The magnitudes of the forces that can cause EIC vary over a wide range and include forces that, under non-corrosive conditions, would be considered negligible. EIC is the most common cause of corrosion in implants for bone applications (Bundy et al., 1983; Lewis et al., 2005) and, because of its localized nature, may go unnoticed until catastrophic failure.

Clinical Relevance of Corrosion

Corrosion of metallic implants, a topic extensively discussed in orthopedic literature, may jeopardize the mechanical stability of the implant and the integrity of the surrounding tissue (Jacobs et al., 1998; Gilbert et al., 2009). Implant failure in the form of aseptic loosening, or osteolysis, may result from metal release in the form of wear debris or electrochemical products generated during corrosion events (Dorr et al., 1990; Jacobs et al., 2001, 2009). Metal ions such as Ti⁴⁺, Co²⁺, and Al³⁺ have been shown to decrease DNA synthesis, mitochondrial dehydrogenase activity, mineralization, and mRNA expression of alkaline phosphatase and osteocalcin in ROS 17/2.8 cells (Sun et al., 1997). Similarly, phagocytosis of Ti particles caused cytotoxicity in a concentration- dependent manner in neonatal rat calvarial osteoblasts (Pioletti et al., 1999) and MG63 cells (Lohmann et al., 2000a).

While implant loosening is less prominent in the dental literature, metal traces originating from dental implants have been found in blood, liver, lungs, and lymph nodes (Lugowski et al., 1991; Smith et al., 1997; Finet et al., 2000). These metal ions and wear debris may also contribute to aseptic loosening by promoting inflammatory complications that may result in macrophage activation, bone resorption, and, rarely, in the potential development of neoplasia (Poggio, 2007; McGuff et al., 2008). Recently, titanium dioxide (TiO₂) was classified as possibly carcinogenic to human beings (i.e., group 2B) at the International Agency for Research on Cancer (IARC) (Baan et al., 2006). Animal studies in rodents provided sufficient evidence of the carcinogenic effects of TiO₂, although epidemiological cohort studies in humans were inconclusive. Furthermore, the immediate and systemic cytotoxic and neoplastic effects of corrosion remain controversial because of conflicting studies that have found no effects of Ti ions or Ti particles on cells (Doran et al., 1998). Moreover, the nanograms of metal per gram of tissue found in vivo (Frisken et al., 2002; Hanawa, 2004) are difficult to compare with the micrograms and milligrams of metal per milliliters of solution used to create an effect in in vitro studies (Pioletti et al., 1999; Lohmann et al., 2000a).

The electrical implications of corrosion and its effect on the surrounding tissue may be an important key to this puzzle, but such effects still remain unclear. Corrosion events generate electrical currents due to electron transfer from ions in the solution to the metallic surface where reactions are occurring. These abnormal currents, and coupled electrical potentials, are directly related to the cyclic loads applied to the implant (Goldberg and Gilbert, 1997; Gilbert et al., 2009). In dental and orthopedic applications, cyclic loads are to be expected from the forces exerted after every bite or every step, respectively. Consequently, it is fair to suggest that cells and tissues in individuals with implants are exposed to abnormal electrical signals for extended periods of time. As described previously, bone cells are sensitive to electrical signals and, thus, could be strongly affected by these corrosion currents. Moreover, these abnormal electrical signals may provide an alternate explanation for the unresolved causes of inflammatory complications and eventual aseptic loosening.

With the growing popularity of treatments like early implant loading, it is imperative to consider the effects of electrical signals on the early stages of osseointegration as well as on long-term outcome. The concern of reducing implant corrosion might be addressed and is being addressed by different methods such as: new formulations of metallic alloys that improve the mechanical and corrosion properties of the implant (Yamazoe et al., 2007; Mareci et al., 2009; Oliveira and Guastaldi, 2009); surface modifications that stabilize the reactivity of the surface (Papakyriacou et al., 2000); or electrical protection (i.e., stimulation) of implants. However, a fundamental understanding of the consequences of abnormal electrical signals on the growth and development of cells and tissues is required for the design of appropriate solutions and adequate treatment for affected individuals.