2.  Requirements for Successful Implant Systems

2.1.  Safety Concerns
Safety concerns should not be limited to dental implants, but also to all dental devices. Specifications and standards have been developed to aid producers, users, and consumers in the evaluation of the safety and effectiveness of dental products. However, the decision of products to test their materials according to national and international standards is purely voluntary [10]. Until the passage in 1976 of the Medical Device Amendments to the Food and Drug Act, medical and dental materials and devices for use in the human were not regulated by any agency of the United States government. The only exception was materials for which therapeutic claims were made, which allowed the Food and Drug Administration (FDA) to consider them as a drug. The Medical Device Amendments of 1976 gave the FDA jurisdiction over all materials, devices, and instruments used in the diagnosis, cure, mitigation, treatment, or prevention of disease in man. This includes materials used professionally and the over-the-counter products sold directly to the public. The Dental Panel places an item one of three classes: Class I, materials posing minimum risk: these are subject only to good manufacturing and recordkeeping procedures. Class II, materials for which safety and efficacy needs to be demonstrated and for which performance standards are available: materials must be shown to meet the performance standard. Class III, materials that pose significant risk and materials for which performance standards have not been formulated: this class is subject to premarket approval by the FDA for safety and efficacy, in much the same manner as a new drug [11].
According to ISO specifications [12], implant devices are required to evaluate several tests; for Group I tests (cytotoxicity tests: ISO 7405, 6.1 and 6.2, and cytotoxicity tests: ISO 10993.5), for Group II tests (subchronic systemic toxicity – Oral application: ISO 10993.11.6.7.1, skin irritation and intracutaneous reactivity: ISO 10993-11.5.2, and sensitization systemic toxicity–application by inhalation: ISO 10993.11.6.7.3, genotoxicity: ISO 10993.3, and local effects after implantation: ISO 10993.6), and for Group III (pulp cappling and pulpotomy: ISO 7405.6.4, and endodontic usage test: ISO 7405.6.5).

2.2.  Compatibility
The implantation of devices for the maintenance or restoration of a body function imposes extraordinary requirements on the materials of construction. Foremost among these is an issue of biocompatibility. There are interactions between the foreign material and the surrounding host living tissue, fluid and blood elements. Some of these are simply adaptive. Others constitute hazard, both short and long term, to the survival of the living system [13,14]. There are mechanical and physical properties which the material must provide, and structural nature which the system should exhibit. Some of these govern the ability of the device to provide its intended function from an engineering viewpoint. Others such as tribology (specifically friction and wear), corrosion and mechanical compliance relate significantly to the biocompatibility concerns. Human implantation applications impose more stringent requirements on reliability than any other engineering task. In most applications an implanted device is expected to function for the life of the patient. As the medical profession becomes more emboldened, the device lifetime must stretch to more than 30 years (if follow-up maintenance is carefully and thoroughly performed and excellent coorperation from patients is obtained), and there are very few engineering devices which have been designed to function for more than 30 years. It is necessary to think in terms of performance reliability of thousands of devices for the lifetime of a patient and a tolerable expectation of failure of perhaps no more than one in one thousand [14,15]. One of many universal requirements of implants, wherever they are used in the body, is the ability to form a suitably stable mechanical unit with the neighboring hard or soft tissues. A loose (or unstable) implant may function less efficiently or cease functioning completely, or it may induce an excessive tissue response. In either case, it may cause the patient discomfort and pain.
There are at three least major required compatibilities for placed implants to exhibit biointegration to receiving hard tissue and biofucntionality thereafter. They include (1) biological compatibility, (2) mechanical compatibility, and (3) morphological compatibility to receiving host tissues [16,17].

2.2.1.  Biological Compatibility (Biocompatibility)
The service conditions in the mouth are hostile, both corrosively and mechanically. All intraorally placed parts are continuously bathed in saliva, an aerated aqueous solution of about 0.1 N chlorides, with varying amounts of Na, K, Ca, PO4, CO2, sulphur compounds and mucin [18]. The pH value is normally in the range of 5.5 to 7.5, but under plaque deposits it can be as low as 2. Temperatures can vary ±36.5 °C, and a variety of food and drink concentrations apply for short periods. Loads may be up to 1,000 N (with normal masticatory force ranging from 150 N to 250 N) [18], sometimes at an impact-load superimposed. Trapped food debris may decompose to create sulphur compounds, causing placed devices discoloration [18]. With such hostile conditions, biocompatibility of metallic materials essentially equates to corrosion resistance because it is thought that alloying elements can only enter the surrounding organic system and develop toxic effects by conversion to ions through chemical or electrochemical process.
As mentioned before, the interface zone between placed implant and receiving hard/soft tissue is strongly governing the success and longevity of placed implant. Hence, there are two major research approaches; one is looking at the interface from foreign implant material side, the other is from vital tissue side. Although this review’s principle scope is aiming at the material’s side, it would be worth to take a brief look at what is happening right after the placing the implant surgically at vital traumatized hard tissue.
After implant placement, initial healing of the bony compartment is characterized by formation of blood clots at the traumatized wound site, protein adsorption and adherence of polymorphonuclear leukocyte [19]. Then approximately two days after placement of the implant, fibroblasts proliferate into the blood clot, organization begins, and an extra-cellular matrix is produced. Approximately one week after the implant is placed, appearance of osteblast-like cells and new bone is seen. New bone reaches the implant surface by osseoconduction (through growth of bone over the surface and migration of bone cells over the implant surfaces) [19].
Why do titanium and its alloys show such good biocompatibility compared with other alloys? The answer to this question is generally that titanium is passive in aqueous solutions, and the passive film that forms on titanium is stable, even in a biological system including chemical and mechanical environments. Such an interpretation is true in many cases. However, the presence of the passive film is only part of the answer when we consider the complex interfacial phenomena to be found between titanium and a biological system, in both biological and biomechanical environments [20,21]. There are certain criteria for any potential metallic materials to be evaluate excellent corrosion resistant, including (1) ease to be oxidized, (2) strong adherence of formed oxide to the substrate, (3) dense of formed oxide, and (4) protectiveness of formed oxide. The Pilling –Bedworth (P-B) ratio is the very simple indication to judge whether the formed oxide is protective or not [22]. If P-B ration is less than 1, since oxide occupies small volume than the metal, so that formed oxide is porous and non-protective. If it is larger than 2, since oxide occupies a large volume and may flake from the surface, exposing fresh substrate surface and again exhibits non-protectiveness. If P-B ration is between 1 and 2, the volume of oxide is similar to that of metal, so that the formed oxide is adherent to substrate, nonporous, and protective. It was calculated that P-B ratio for TiO2 formation is 1.76, indicating that the formed TiO2 is protective. Titanium is a highly reactive metal and will react within microseconds to form an oxide layer when exposed to the atmosphere [23]. Although the standard electrode potential was reported in a range from −1.2 to −2.0 volts for the Ti ↔ Ti3+ electrode reaction [24], due to strong chemical affinity to oxygen, it easily produces a compact oxide film, ensuring high corrosion resistance of the metal. This oxide, which is primarily TiO2, forms readily because it has one of the highest heats of reaction known (ΔH = −915 kJ/mole) (for 298.16°–2,000°K) [25]. It is also quite impenetrable to oxygen (since the atomic diameter of Ti is 0.29 nm, the primary protecting layer is only about 5 to 20 atoms thick) [26]. The formed oxide layer adheres strongly to the titanium substrate surface. The average single-bond strength of the TiO2 to Ti substrate was reported to be about 300 kcal/mol, while it is 180 kcal/mol for Cr2O3/Cr, 320 kcal/mol for Al2O3/Al, and 420 kcal/mol for both Ta2O5/Ta and Nb2O5/Nb [27].
During implantation, titanium releases corrosion products (which is mainly titanium oxide or titanium hydro-oxide) into the surrounding tissue and fluids even though it is covered by a thermodynamically stable oxide film [28,29]. An increase in oxide thickness, as well as incorporation of elements from the extra- cellular fluid (P, Ca, and S) into the oxide, has been observed as a function of implantation time [30]. Moreover, changes in the oxide stoichiometry, composition, and thickness have been associated with the release of titanium corrosion products in vitro [31]. Properties of the oxide, such as stoichiometry, defect density, crystal structure and orientation, surface defects, and impurities were suggested as factors determining biological performance [32,33].
The performance of titanium and its alloys in surgical implant applications can be evaluated with respect to their biocompatibility and capability to withstand the corrosive species involved in fluids within the human body [34]. This may be considered as an electrolyte in an electrochemical reaction. It is well documented that the excellent corrosion resistance of titanium materials is due to the formation of a dense, protective, and strongly-adhered film – which is called a passive film, as discussed before. Such a surface situation is referred to passivity or a passivation state. The exact composition and structure of the passive film covering titanium and its alloys is controversial. This is the case not only for the “natural” air oxide, but also for films formed during exposure to various solutions, as well as those formed anodically. The “natural” oxide film on titanium ranges in thickness from 2 to 7 nm, depending on such parameters as the composition of the metal and surrounding medium, the maximum temperature reached during the working of the metal, the surface finish, etc.
Oxides formed on Ti materials are varied with a general form; TiOX (1 < x < 2). Depending on x values, there are five different crystalline oxides; i.e., (1) cubic TiO (ao = 4.24 Å), (2) hexagonal Ti2O3 (ao = 5.37 Å, α = 56°48’), (3) tetragonal TiO2 (anatase) (ao = 3.78Å, co = 9.50 Å), (4) tetragonal TiO2 (rutile) (ao = 4.58 Å, co = 2.98 Å), and (5) orthorhombic TiO2 (brookite) (ao = 9.17 Å, bo = 5.43 Å, co = 5.13 Å). Besides these, there are (6) non-stoichiometric oxide (when x is not integral), and (7) amorphous oxides. It is widely believed that, among these oxides, only rutile and anatase type oxides are stable at normal conditions. Of interest, choice for rutile formation or anatase formation depends on the acidity of used electrolyte [8]. The rutile and anatase type oxides exibit different physical properties – interms of surface tension. Lim et al. [35] prepared various surface conditions on pure titanim and measured surface contact angles, surface electrochemical potential and roughness. It was found that the surface covered with only rutile type TiO2 was hydrophobic, whereas that covered with a mixture of rutile and anatase type of oxides showed hydrophilicity [35].
The level of neutrophil priming and activation following implant placement may be linked to the development and maintenance of long-term stability and osseointegration. Bisphosphonate effect on neutrophil activation was examined on differently treated surfaces [36]. Neutrophils were isolated from whole blood collected from healthy human donors, on a double dextran gradient. Treated surfaces were incubated with 5 × 105 neutrophils per curette. Luminol-dependent CL (chemiluminescence) was recorded for 60 min (priming or inflammatory phase), followed by secondary stimulation with 10−7 M phorbol myrisitate acetate at 60 min (activation phase) for a continuous CL measurement over 120 min. SEM evaluation was preformed. Results indicated that titanium surfaces which were covered with a mixture of rutile and anatase type TiO2 oxide films are capable of priming neutrophils, when compared to the acid-treated surface which was covered with rutile oxide only [36].
Using Auger Electron Spectroscopy (AES) to study the change in the composition of the titanium surface during implantation in human bone, observed that the oxide formed on titanium implants grows and takes up minerals during the implantation [30,37]. The growth and uptake occur even though the adsorbed layer of protein is present on the oxide, indicating that mineral ions pass through the adsorbed protein. It was shown that, using Fourier Transform Infrared Reflection Absorption Spectroscopy (FTIR-RAS), phosphate ions are adsorbed by the titanium surface after the protein has been adsorbed. Using x-ray photoelectron spectroscopy (XPS) [38], it was demonstrated that oxides on commercially pure titanium and titanium alloy (Ti-6Al-4V) change into complex phosphates of titanium and calcium containing hydroxyl groups which bind water on immersion in artificial saliva (pH = 5.2) [39]. It was shown that titanium is in almost direct contact to bone tissue, separated only by an extremely thin cell-free non-calcified tissue layer. Transmission electron microscopy revealed an interfacial hierarchy that consisted of a 20–40 nm thick proteoglycan layer within 4 nm of the titanium oxide, followed by collagen bundles as close as 100 nm and Ca deposits within 5 nm of the surface [40]. To reach the steady-state interface described, both the oxide on titanium and the adjacent tissue undergo various reactions. The physiochemical properties of titanium have been associated with the unique tissue response to the materials: these include the biochemistry of released corrosion products, kinetics of release and the oxide stoichiometry, crystal defect density, thickness and surface chemistry [41]. All these studies indicate that the surface oxide on titanium materials reacts with mineral ions, water, and other constituents of biofluids, and that these reactions, in turn, cause a remodeling of the surface.
As seen in the above, in general, the titanium passivating layer not only produces good corrosion resistance, but it seems also to allow physiological fluids, proteins, and hard and soft tissue to come very close and/or deposit on it directly. Reasons for this are still largely unknown, but it may have something to do with things such as the high dielectric constant for TiO2 (50 to 170 vs. 4–10 for alumina and dental porcelain), which should result in considerably stronger van der Waal’s bonds on TiO2 than other oxides; TiO2 may be catalytically active for a number of organic and inorganic chemical interactions influencing biological processes at the implant interface. The TiO2 oxide film may permit a compatible layer of biomolecule to attach [42,43].

2.2.2.  Mechanical Compatibility
Biomechanics involved in implantology should include at least (1) the nature of the biting forces on the implants, (2) transferring of the biting forces to the interfacial tissues, and (3) the interfacial tissues reaction, biologically, to stress transfer conditions. Interfacial stress transfer and interfacial biology represent more difficult, interrelated problems. While many engineering studies have shown that variables such as implant shape, elastic modulus, extent of bonding between implant and bone etc., can affect the stress transfer conditions, the unresolved question is whether there is any biological significance to such differences. The successful clinical results achieved with osseointegrated dental implants underscore the fact that such implants easily withstand considerable masticatory loads. In fact, it was reported that bite forces in patients with these implants were comparable to those in patients with natural dentitions. A critical aspect affecting the success or failure of an implant is the manner in which mechanical stresses are transferred from the implant to bone smoothly [44]. It is essential that neither implant nor bone be stressed beyond the long-term fatigue capacity. It is also necessary to avoid any relative motion that can produce abrasion of the bone or progressive loosening of the implants. An osseointegrated implant provides a direct and relatively rigid connection of the implant to the bone. This is an advantage because it provides a durable interface without any substantial change in form or duration. There is a mismatch of the mechanical properties at the interface of Ti and bone. It is interesting to observe that from a mechanical standpoint, the shock-absorbing action would be the same if the soft layer were between the metal implant and the bone. In the natural tooth, the periodontum, which forms a shock-absorbing layer, is in this position between the tooth and jaw bone [44]. Natural teeth and implants have different force transmission characteristics to bone. Compressive strains were induced around natural teeth and implants as a result of static axial loading, whereas combinations of compressive and tensile strains were observed during lateral dynamic loading [8,9]. Magnitude of strain around the natural tooth is significantly lower than the opposing implant and occluding implants in the contra lateral side for most regions under all loading conditions. It was repoted that there was a general tendency for increased strains around the implant opposing natural tooth under higher loads and particularly under lateral dynamic loads [45].
By means of finite element (FEM) analysis, stress-distribution in bone around implants was calculated with and without stress-absorbing element [46]. A freestanding implant and an implant connected with a natural tooth were simulated. For the freestanding implant, it was reported that the variation in the modulus of elasticity of the stress-absorbing element had no effect on the stresses in bone. Changing the shape of the stress-absorbing element had little effect on the stresses in cortical bone. For the implant connected with a natural tooth, it was concluded that (i) a more uniform stress was obtained around the implant with a low modulus of elasticity of the stress-absorbing element, and (ii) the bone surrounding the natural tooth showed a decrease in the height of the peak stresses.
The dental or orthopedic prostheses, particularly the surface zone thereof, should respond to the loading transmitting function. The placed implant and receiving tissues establish a unique stress-strain field. Between them, there should be an interfacial layer. During the loading with implant/bone couple, the strain-field continuity should be held (if not, it should indicate that implant is not fused to vital bone), although the stress-field is obviously in a discrete manner due to different values of modulus of elasticity (E) between host tissue and foreign implant material. Namely, stress at bone σB = EBɛB and stress at implant σI = EIɛI. Under the continuous strain field, ɛB = ɛI. However EB ≠ EI due to dissimilar material couple condition. If the magnitude of the difference in modulus of elasticity is large, then the interfacial stress, accordingly, could become so large that the placed implant system will face a risky failure or detachment situation. In other words, if interfacial stress due to stress difference Δσ = (σI–σB) is larger than the osteointegrated fused implant retention strength, the placed impant will be failed. Therefore, materials for implant or surface zone of implants should be mechanically compatible to mechanical properties (especially, modulus of elasticity) of receiving tissues to minimize the interfacial discrete stress. This is the second compatibility and is called as the mechanical compatibility [8,16].
Figure 1 shows a relationship between yield strength and modulus of elasticity (in other words, rigidity) of various types of biomaterials and bones which are receiving vital tissue for placed implants. As seen clearly, strength and rigidity of all biomaterials concerned are related linearly on both-log scale. From the point of strain-continuity viewpoint, it is ideal to choose any implantable materials that have both strength and rigidity values close to those of receiving bone. Hydroxyapatite coating onto titanium implant has been widely adopted since the both hydroxyapatite (HA) and receiving vital bone possess similar chemical compositions, hence early adaptation can be highly expected. At the same time, EHA is positioned inbetween the values of EB and EI; as a result, HA coating will have a second function for mechanical compatibility to make the stress a smooth transfer (or to minimize the interfacial stress). This is one of the typical hindsight, because HA-coating is originally and still now performing due to its similarity of its chemical composition to receiving bone.

2.2.3.  Morphological Compatibility
Surface plays a crucial role in biological interactions for four reasons. First, the surface of a biomaterial is the only part in contact with the bioenvironment. Second, the surface region of a biomaterial is almost always different in morphology and composition from the bulk. Differences arise from molecular rearrangement, surface reaction, and contamination. Third, for biomaterials that do not release nor leak biologically active or toxic substances, the characteristics of the surface govern the biological response. And fourth, some surface properties, such as topography, affect the mechanical stability of the implant/tissue interface [47,48].
In a scientific article [16], it was found that surface morphology of successful implants has an upper and lower limitations in average roughness (1–50 μm) and average particle size (10–500 μm), regardless of types of implant materials (either metallic, ceramics, or polymeric materials). If a particle size is smaller than 10 μm, the surface will be more toxic to fibroblastic cells and have an adverse influence on cells due to their physical presence independent of any chemical toxic effects. If the pore is larger than 500 μm, the surface zone does not maintain sufficient structural integrity because it is too coarse. This is the third compatibility – morphological compatibility [16,17].
It has been shown that preparation methods of implant surface can significantly affect the resultant properties of the surface and subsequently the biological responses that occur at the surface [49–51]. Recent efforts have shown that the success or failure of dental implants can be related not only to the chemical properties of the implant surface, but also its macromorphologic nature [52–55]. From an in vitro standpoint, the response of cells and tissues at implant interfaces can be affected by surface topography or geometry on a macroscopic basis [53,55], as well as by surface morphology or roughness on a microscopic level [53,56]. These characteristics undoubtly affect how cells and tissues respond to various types of biomaterials. Of all the cellular responses, it has been suggested that cellular adhesion is considered the most important response necessary for developing a rigid structural and functional integrity at the bone/implant interface [57]. Cellular adhesion alters the entire tissue response to biomaterials [58].
The effect of surface roughness (Ra: 0.320, 0.490, and 0.874 μm) of the titanium alloy Ti-6Al-4V on the short- and long-term response of human bone marrow cells in vitro and on protein adsorption was investigated [59]. Cell attachment, cell proliferation, and differentiation (alkaline phosphatase specific activity) were determined. The protein adsorption of bovine serum albumin and fibronectin, from single protein solutions on rough and smooth Ti-6Al-4V surfaces was examined with XPS and radio labeling. It was found that (i) cell attachment and proliferation were surface roughness sensitive, and increased as the roughness of Ti-6Al-4V increased, (ii) human albumin was adsorbed preferentially onto the smooth substratum, and (iii) the rough substratum bound a higher amount of total protein (from culture medium supplied with 15% serum) and fibronectin (10-fold) than did the smooth one [59], suggesting an importance of the rugophilicity.
Events leading to integration of an implant into bone, and hence determining the long-time performance of the device, take place largely at the interface formed between the tissue and the implant [60]. The development of this interface is complex and is influenced by numerous factors, including surface chemistry and surface topography of the foreign material [61–65]. For example, Oshida et al. treated NiTi by acid-pickling in HF-HNO3-H2O (1:1:5 by volume) at room temperature for 30 seconds to control the surface topology and selectively dissolve Ni, resulting in a Ti-enriched surface layer [66], demonstrating that surface topology can be easily controlled.
The role of surface roughness on the interaction of cells with titanium model surfaces of well-defined topography was investigated using human bone-derived cells (MG63 cells). The early phase of interactions was studied using a kinetic morphological analysis of adhesion, spreading, and proliferation of the cells. SEM and double immuno-fluorescent labeling of vinculin and actin revealed that the cells responded to nanoscale roughness with a higher cell thickness and a delayed apparition of the focal contacts. A singular behavior was observed on nanoporous oxide surfaces, where the cells were more spread and displayed longer and more numerous filopods. On electrochemically micro-structured surfaces, the MG63 cells were able to penetrate inside, adhere, and proliferate in cavities of 30 or 100 μm in diameter, whereas they did not recognize the 10 μm diameter cavities. Cells adopted a 3D shape when attaching inside the 30 μm diameter cavities. It was concluded that nanotopography on surfaces with 30 μm diameter cavities had little effect on cell morphology compared to flat surfaces with the same nanostructure, but cell proliferation exhibited a marked synergistic effect of microscale and nanoscale topography [67].
On a macroscopic level (roughness > 10 μm) roughness influences the mechanical properties of the titanium/bone interface, the mechanical interlocking of the interface, and the biocompatibility of the material [68,69]. Surface roughness in the range from 10 nm to 10 μm may also influence the interfacial biology, since it is the same order as the size of the cells and large biomolecules [23]. Microroughness at this level includes material defects, such as grain boundaries, dislocation steps and kinks, and vacancies that are active sites for adsorption, and therefore influence the bonding of biomolecules to the implant surface [70]. Microrough surfaces promote significantly better bone apposition than smooth surfaces, resulting in a higher percentage of bone in contact with the implant. Microrough surfaces may influence the mechanical properties of the interface, stress distribution, and bone remodeling [71]. Increased contact area and mechanical inter-locking of bone to a microrough surface can decrease stress concentrations resulting in decreased bone resorption. Bone resorption takes place shortly after loading smooth surfaced implants [72], resulting in a fibrous connective tissue layer, whereas remodeling occurs on rough surfaces [73].
Recently developed clinical oral implants have been focused on topographical changes of implant surfaces, rather than alterations of chemical properties [55,74–77]. These attempts may have been based on the concept that mechanical interlocking between tissue and implant materials relies on surface irregularities in the nanometer to micron level. Recently published in vivo investigations have shown significantly improved bone tissue reactions by modification of the surface oxide properties of Ti implants [78–85]. It was found that in animal studies, bone tissue reactions were strongly reinforced with oxidized titanium implants, characterized by a titanium oxide layer thicker than 600 nm, a porous surface structure, and an anatase type of Ti oxide with large surface roughness compared with turned implants [83,84]. This was later supported by work done by Lim et al. [35], Oshida and [86], and Elias et al. [87] who found that the alkali-treated CpTi surface was covered mainly with anatase type TiO2, and exhibited hydrophilicity, whereas the acid-treated CpTi was covered with rutile type TiO2 with hydrophobicity. Besides this characteristic crystalline structure of TiO2, it was mentioned that good osseointegration, bony apposition, and cell attachment of Ti implant systems [28,88,89] are partially due to the fact that the oxide layer, with unusually high dielectric constant of 50–170, depending on the TiO2 concentration, may be the responsible feature [23,42,43].

2.3.  MRI Safety and Image Compatibility
Magnetic resonance imaging (MRI) is a technology developed in medical imaging that is probably the most innovative and revolutionary other than computed tomography. MR is a three-dimensional imaging technique used to image the protons of the body by employing magnetic fields, radio frequencies, electromagnetic detectors, and computers [90]. For millions of patients worldwide, MRI examinations provide essential and potentially life-saving information. Some devices, such as pacemakers and neurostimulators, have limitations related to MRI safety and may be contraindicated for use with MRI. Even more devices, such as stents, vena cava filters, and some types of catheters and guidewires, are safe for use with MRI but have limited MRI image compatibility. Some of these devices are simply not well-imaged under MRI. Others have properties that interfere with the MRI image by causing an image artifact (distortion) in the area in and around the device, limiting the effectiveness of MRI for assisting placement or diagnostic follow up on these implants. It may be contraindicated in certain situations because the magnetic field present in the MRI environment may, under certain circumstances, result in movement or heating of a metallic orthopaedic implant device. Metals that exhibit magnetic attraction in the MRI setting may be subject to movement (deflection) during the procedure. Both magnetic and non-magnetic metallic devices of certain geometries may also be subjected to heating caused by interactions with the magnetic field. Of secondary concern, is the possibility of image artifacts that can compromise the procedure and image quality.
There are currently several researchers as well as an ASTM committee exploring methods for accurately assessing the MRI compatibility of implant devices. The primary focus of the research has been the measurement of implant movement in response to a magnetic field. Shellock and co-workers [91–93] conducted several studies in which the movement/deflection of various orthopaedic implants was measured in the high magnetic field (0.3–1.5 Tesla) region of MRI units. The results of these studies show no measurable movement of implants fabricated from cobalt, titanium and stainless steel alloys. The movemen/deflection of selected orthopaedic implants in a 3.0 Tesla MRI unit was also examined and it was found that devices fabricated from cobalt, titanium and stainless steel exhibited little or no movement/deflection [94].
Ferromagnetic metal will cause a magnetic field inhomogeneity, which, in turn, causes a local signal void, often accompanied by an area of high signal intensity, as well as a distortion of the image. They create their own magnetic field and dramatically alter precession frequencies of protons in the adjacent tissues. Tissues adjacent to ferromagnetic components become influenced by the induced magnetic field of the metal hardware rather than the parent field and, therefore, either fail to precess or do so at a different frequency and hence do not generate useful signal. Two components contribute to susceptibility artifact, induced magnetism in the ferromagnetic component itself and induced magnetism in protons adjacent to the component. Artifacts from metal may have varied appearances on MRI scans due to different type of metal or configuration of the piece of metal. In relation to imaging titanium alloys are less ferro-magnetic than both cobalt and stainless steel, induce less susceptibility artifact and result in less marked image degradation [94–96].