Читать книгу Fundamentals of Implant Dentistry - John III Beumer - Страница 14

CHAPTER 4

Contemporary Implant Materials

Neil Waddell | Kai Chun Li | Abdullah Barazanchi | Kumar C. Shah Basil Al-Amleh | Karl M. Lyons | Benjamin M. Wu

The previous chapter provided the fundamental biomechanics rationale for prosthetic design to manage occlusal forces and moments. This chapter continues the discussion by focusing on the materials that provide resistance to occlusal loads. The ideal material should provide long-term function specifically for the intended purpose without altering the structure and function of surrounding natural tissues and other synthetic materials and without being altered by the chemical, biochemical, mechanical, and biologic factors in the local microenvironment. Because all materials break down by degradation, erosion, corrosion, and wear, the ideal material should also produce minimal breakdown products, and the debris should be biocompatible with the local and systemic physiology and immune systems. Therefore, materials not only directly determine the resistance to the applied stress and strain and govern the stress distribution to the supporting bone and soft tissues; they also influence local microbiology and tissue response. The following sections describe the materials of each critical component within the implant-prosthesis system (implant fixtures, prosthesis, substructure, bars, and teeth) to highlight the important role that each material has on long-term clinical success. Detailed metallurgy of titanium is beyond the scope of this chapter, but relevant materials science is discussed briefly where appropriate to explain the roles that material composition and processing play in final material properties.

Implant Materials

Titanium

In the search for the ideal implant fixture material, specific properties are considered desirable. For implant fixtures, the material used should be biocompatible and resistant to wear and corrosion, and should possess high yield strength, high fatigue strength, and fracture resistance yet ideally have elastic properties similar to bone. Several material options have been trialed for use, including metals, ceramics, and polymers. Among all materials investigated, titanium has remained the material of choice for over 30 years. The material’s biocompatibility and resistance to corrosion promote stable osseo- integration with the surrounding bone while also fulfilling most of the required physical properties at a relatively low cost in terms of manufacturing and post-processing.

There are 38 grades of titanium, and several grades are commonly used in dentistry. The grades differ in elemental composition that control the microstructure, corrosion resistance, ductility, strength, and suitability as dental implant materials. Grades 1 to 4, commonly called commercially pure (CP) titanium, are unalloyed with trace elements (0.2–0.5 wt% Fe, 0.18–0.4 wt% Ti) that strengthen the CP titanium. Grade 1, with the lowest iron and oxygen content, is the softest CP titanium. Grade 4, with the highest iron and oxygen content, is the strongest pure titanium and the most commonly used grade for dental implant fixtures. In general, physical properties such as elastic modulus, tensile strength, and yield strength increase with higher titanium grades (Grade 4 is strongest), while ductility and percent elongation decrease (Grade 1 is most ductile). Grade 5 is titanium alloyed with aluminum (Al) and vanadium (V), or Ti-6Al-4V. The most commonly used titanium alloy, Ti-6Al-4V is much stronger than CP titanium. Grade 23, sometimes referred to as surgical titanium, is Ti-6Al-4V with extra low interstitials (ELI). By reducing the interstitial elements iron and oxygen during the processing, Ti-6Al-4V ELI enhances compositional purity and biocompatibility and optimizes strength, ductility, and fracture toughness. The role of aluminum, vanadium, and interstitial elements are discussed in the next section.

Pure titanium contains two stable phases: (1) low- temperature α phase with hexagoanal close-packed (HCP) crystalline structure and (2) high-temperature β phase with body-centered cubic (BCC) crystalline structure above the transus temperature of 882.5°C. In general, HCP crystals have three slip systems, as this crystal structure offers one plane with the greatest atomic density within the unit cell and three directions that have the highest atomic density along the slip plane. Dislocation motion along these three slip systems is preferred for plastic deformation, or ductility. In contrast, BCC has 48 slip systems (12 planes with 2 directions each, plus 6 planes with 4 directions each, 12 planes × 2 directions + 6 planes × 4 directions = 48 systems), and the abundance of slip systems allow a higher degree of freedom for dislocation motion, hence much higher ductility. In general, β titanium tends to be more ductile, weaker than α titanium, and metastable, as described later. CP titanium (Grades 1 to 4) are α alloys and exhibit enough ductility (24% elongation for Grade 1 to 15% elongation for Grade 4) for dental implant applications (Table 4-1).

The transformation of the BCC β to the HCP α pure titanium is achieved by cooling slowly below the transformation temperature (transus). If BCC β is cooled rapidly (quenched), the material will retain the high temperature β phase with local metamorphic changes into metastable structures. Substitutional and interstitial impurities can significantly alter the transformation temperature and physical properties of titanium. Substitutional impurity atoms can replace titanium atoms in the crystal lattice, while interstitial impurity atoms can occupy the small spaces between adjacent titanium atoms. Since α is the lower temperature phase, impurities that can increase the transus temperature are considered α stabilizers, because they delay the transition to β phase at temperatures above the pure titanium transus temperature of 882.5°C. Examples of α stabilizers include substitutional aluminum and interstitial elements nitrogen, oxygen, and carbon, which tend to have higher solubility in the α phase than β phase. Aluminum is a popular additive for biomedical titanium alloys (eg, Ti-6Al-4V) due to its biocompatibility, low density, strengthening effects on titanium, and high solubility limit of about 8 wt% in titanium.

In contrast, impurities that have higher solubility in β phase will decrease the transus temperature, thereby stabilizing the β phase at lower temperature than 882.5°C. β stabilizers are divided into two categories. β-isomorphous substitutional elements such as vanadium, tantalum, molybdenum, niobium, and the interstitial element hydrogen lower the transus temperature more effectively with increasing impurity concentration up to their solubility limits. Vanadium is the most commonly used β stabilizer in biomedical implant grade titanium. Grade 5 (Ti-6Al-4V) and Grade 23 (Ti-6Al-4V ELI) titanium are comprised of both α+β alloys and intermetallic compounds, with aluminum stabilizing the α alloy while vanadium stabilizes the β alloy at room temperature. Both maintain the corrosion resistance of titanium yet significantly improve the alloy’s strength. Grade 5 titanium has almost twice the ultimate tensile strength and yield strength of Grade 4 titanium.

Grade 5 and Grade 23 alloys are the workhorses of biomedical titanium implants, but other isomorphous (β) stabilizers such as Mo and Nb can replace 5. For example, Ti-6Al-7Nb alloys are used in commercial hip implants when biocompatibility to 5 is a concern.

β-eutectoid stabilizers such as chromium, iron, manganese, silicon, and copper have lower solubility in β titanium and require lower critical concentration than the isomorphous β stabilizers to retain β phase after quenching. Eutectoid stabilizers tend to form intermetallic compounds by eutectoid decomposition of the β phase into two new phases (β α + Ti_xE_y), where Ti_xE_y is the equilibrium intermetallic compound formed by Ti and the eutectoid element. Note that both (α) and Ti_xE_y compound have different composition from the original (β) titanium. While α titanium cannot be strongly influenced by heat treatment, eutectoid decomposition effectively facilitates significant strengthening by controlled decomposition of β phase to α phase by a simple heat treatment.

A third group of impurities are neutral stabilizers that are soluble in both α and β phases and strengthen both phases without changing the transus temperature or favoring either phase (hence “neutral”). Examples of neutral stabilizers include zirconium and tin. Zirconium has complete solid solubility in titanium and very similar chemical and physical properties to titanium. As discussed later, titanium-zirconium dental implants are currently available commercially.

Although Grade 4 titanium has satisfactory mechanical properties when used in the fabrication of regular-diameter implants, there are concerns regarding its physical properties when used in narrow-diameter (ie, 3- to 3.5-mm) implants. Narrow-diameter implants allow for placement of fixtures in smaller interdental spaces and/or narrower widths of bone without the need for additional bone grafting procedures, thereby reducing the cost and complexity of procedures. However, with the reduced amount of metal surrounding the prosthetic connection, fatigue resistance and strength of the fixture material become a critical factor in implant success and avoiding fracture of the fixture body. For this purpose, Grade 5 titanium alloys (Ti-6Al-4V) have been used to manufacture narrow-diameter implants due to their superior mechanical properties. However, there are concerns over the use of Grade 5 titanium in orthopedic implants in which significant implant wear can cause the release of vanadium and aluminum into body tissues and cause cell toxicity and type IV hypersensitivity reactions.1 Although soluble vanadium and aluminum have been shown to exhibit cytotoxic effects in vitro, dental implants are not subjected to the type of loading and wear motion that hip implants undergo. While in vitro studies of solubilized α and β stabilizers may not reflect the conditions that dental implant fixtures undergo in the oral cavity, the risk of hypersensitivity reactions to corrosion mechanisms (galvanic, pitting, fatigue, abrasive, and fretting) should be considered when weighing the biologic risk against the enhanced mechanical properties.

More recently, a titanium alloy formulation has been introduced for use in implant fixture fabrication. The new alloy is composed of about 85% titanium and 15% zirconia (Ti-15Zr) and is marketed under the brand Roxolid (Straumann).2 Titanium-zirconium is claimed to have similar or superior strength to Grade 5 titanium (Ti-6Al-4V) and similar corrosion resistance to Grade 4 CP titanium but with higher tensile strength and fatigue strength (40% and 60%, respectively). However, little data exists that directly compares the mechanical properties of titanium-zirconium with various types of titanium, particularly those produced using different manufacturing methods. Also, despite the improved mechanical properties reported with this material, there have still been fixture body fractures reported when using this alloy formulation.3 Clinically, their success at 3 years is comparable to that of regular-diameter implants, although most studies on implant fixtures fabricated using this alloy combination are short term and about narrow- diameter connections.4

Titanium is highly reactive with oxygen and nitrogen, and readily oxidizes in air to form stable titanium dioxide (TiO₂) on the implant surface. In fact, it is this strong reactivity with oxygen and nitrogen that delayed the production of high purity titanium by about 150 years after its initial discovery in the late 1700s. Titanium is extracted from naturally occurring minerals ilmenite (FeTiO₃) and rutile sand (TiO₂). TiO₂ is most commonly arranged in a tetragonal crystal structure, and the two most thermodynamically stable forms at room temperature and pressure for TiO₂ are anatase and rutile. They have similar mechanical properties, but their crystal lattice dimensional differences contribute to significant difference in their interactions with electrons and electromagnetic radiation. For example, rutile can more readily absorb ultraviolet light due to a smaller electron band gap structure and would be more sensitive to ultraviolet light during photofunctionalization (see chapter 2, section entitled “Biologic Aging and Photofunctionalization of Implants”). Cubic, monoclinic, and orthorhombic structures of TiO₂ can be found at higher temperatures or pressures and can be stabilized with dopants to maintain them at physiologic conditions. Depending on the manufacturing conditions (temperature and pressure) and the concentration of dopants or impurities, various mixtures of crystalline structures and amorphous TiO₂ can be produced on implant surfaces. Regardless of the final TiO₂ structure, surface TiO₂ increases resistance against acids, bases, and galvanic corrosion over that of nonoxidized titanium, and TiO₂ is responsible for the biologic activities of titanium.

Due to its prevalent use in implant dentistry, titanium has been reviewed extensively over the past three decades, and a comprehensive discussion on titanium and recent advances in titanium alloy and nanoengineering will not be repeated here. Despite these advances, all titanium alloys suitable for implant fixtures have much higher elastic modulus than bone, and the modulus mismatch has been attributed to the decrease of bone deformation and osteoblast stimulation (stress shielding). Finally, even with the latest anodization methods, titanium is a poor material in terms of esthetics due to its silver-gray color, which is especially problematic when implant fixtures are exposed after soft tissue recession. Along with potential for type IV hypersensitivity reactions in susceptible patients, alternatives to titanium were developed, and some are presented in the following sections. It should be emphasized that none of the alternative implant materials come close to matching the fracture toughness of titanium alloys, which can reach 50 MPa • m½. The fracture toughness of alumina and zirconia generally range from 1 to 10 MPa • m½, but they are highly sensitive to impurity concentration, processing history, and environmental conditions.

Alumina

The first commercially available ceramic implants were primarily aluminum oxide–based ceramics. In the 1960s and 1970s, aluminous porcelains and high-alumina ceramics were developed for the fabrication of all-ceramic crowns.5 Two types of ceramic implants were also developed. One was composed of polycrystalline aluminum oxide and was marketed as the Cerasand implant (Incermed)6 or the Tübingen implants (Friadent).7 The other ceramic implant was a single-crystal aluminum oxide (synthetic sapphire) and was marketed as the Bioceram implant (Kyocera).8 Clinical trials using alumina-based implants are limited, and concerns with implant fracture due to low fracture toughness < 5 MPa • m1/2 as well as varied long-term success resulted in these dental implants being withdrawn from the market.

Zirconia

Well before its popularity in dentistry, zirconia was considered for use in biomedical implants in the late 1960s. Eventually it was introduced in the femoral heads of total hip replacements in an effort to solve the brittleness problem of alumina ball heads. However, due to a series of failures that were reported in 2001, the use of zirconia in orthopedic surgery dropped drastically and became a controversial issue in orthopedics.9,10 Around the same time that zirconia femoral heads were spontaneously failing in total hip replacement prostheses, zirconia was becoming the latest material that dental companies were marketing. This resulted in zirconia being used in a number of restorative applications such as root canal posts,11 frameworks for crowns and fixed partial dentures,12,13 implant abutments,14 and dental implants.15

Located immediately below titanium on the periodic table, zirconium (Zr) forms a lustrous, grayish metal that resembles titanium. Like titanium, zirconium reacts readily with oxygen to form a high-strength, white crystalline zirconium dioxide (ZrO₂), or zirconia. Like TiO₂ and most other oxides, the crystal structure of zirconia depends on its composition and temperature. Its monoclinic crystal structure is the most thermodynamically stable zirconia at room temperature and pressure. Other structures such as cubic crystals can be stabilized with the addition of dopants such as yttria (Y₂O₃). In nature, zirconia is found in the mineral baddeleyite, which has the same tetragonal oxide group structure as rutile (TiO₂). Therefore, extraction of ZrO₂ requires removal of rutile and other contaminants from baddeleyite. Other natural sources of zirconium include zircon sand (ZrSiO₄), from which removal of silica and impurities yields ZrO₂. While zirconium has slightly better durability and corrosion resistance, is less allergenic, and has low toxicity, titanium offers similar mechanical properties at much lower density. Zirconia materials are similar to TiO₂ in strength, corrosion resistance, abrasive wear, and mechanical properties.

Phase transformation

Pure zirconia ceramic exists as a monoclinic crystalline structure at room temperature and pressure. Upon heating, monoclinic zirconia transitions to tetragonal (980°C) and cubic crystals at higher temperatures (2,370°C). Since the tetragonal and cubic forms are unstable at room temperature, the spontaneous phase transformation to the monoclinic phase involves a significant volume expansion that induces large stresses that can crack zirconia during cooling. To avoid the volume expansion, phase transformation is retarded by adding oxide dopants to stabilize the high temperature tetragonal and cubic structures. The most popular stabilizing addition to dental zirconia is 3 to 5 wt% yttria (3 mol% yttria- stabilized tetragonal zirconia polycrystal [3Y-TZP]). Other additives can also improve the durability of dental zirconia. Alumina can dramatically increase the strength of zirconia and is commonly added for ZrO₂ implants and abutments. For restorations where translucency and esthetics are important, the amount of alumina must be kept relatively small, but they tend to be weaker than the opaque materials.

A peculiarity of 3Y-TZP is its so-called “transformation toughening.” This phenomenon occurs when each grain of metastable TZP zirconia crystal is retained in its tetragonal structure at room temperature so that each grain can undergo transformation instead of only the precipitates. Upon clinical loading, local stresses that are high enough to propagate microcracks will also provide the necessary activation energy for grains at the crack tip to undergo tetragonal/monoclinic phase transformation. Because the rest of the material does not undergo phase transformation, the crack tip is constrained against the bulk material, resulting in volume expansion that compresses on the crack tip to retard its growth and effectively enhances the fracture toughness of the material to 10 to 18 MPa • m½, depending on the dopant concentration and processing conditions. Although still far below the fracture toughness of titanium, the use of mestastable tetragonal zirconia has the potential to improve the performance outcomes of dental restorations.

In practice, zirconia is available to dentistry in microporous presintered ceramic blocks and disks. The properties of the ceramic blocks are influenced both by the chosen raw material and the actual press-forming processes that are employed to manufacture zirconia blocks. First, an organic binder is added to the ceramic powder prior to pressing. The powder reduces internal friction during the pressing procedure. Furthermore, the binder gives the pressed block, also referred to as a green body, a certain inherent solidity. Next, the powder is pressed into blocks by a variety of different techniques that may affect its properties, especially during milling and sintering. A partially sintered cold isostatic–pressed zirconia block has the advantage that its chalklike consistency makes it far simpler and faster to process without damaging the material than is the case for a block processed by hot isostatic pressing (HIP). Depending on its density, the partially sintered ZrO₂ undergoes a linear sintering shrinkage of 20% to 25% during full sintering. This corresponds to a volume shrinkage of 49% to 57%. Dimensions of the framework must be linearly increased prior to milling the presintered blocks to compensate for the sintering shrinkage in accordance with the manufacturer’s specifications.

The initial grain size and pores within the presintered blocks can also determine the quality of the sintered zirconia. These are governed by the heating rates, hold times, and the final temperature selected during the sintering process. For example, temperatures that are too high or sintering times that are too long result in so-called “large grain growth.” Individual grains can grow to 100–1,000 times their original volume (0.3–0.4 µm) at the expense of other grains. This results in extremely inhomogeneous grain size distributions with the inherent danger that the metastable tetragonal phase will spontaneously change into the monoclinic phase. The heating rate greatly influences the distribution of pores in the microstructure. Areas with greatly densified ZrO₂ grains and areas with loosely packed ZrO₂ grains result if the heating rate is not precisely controlled. If the sintering temperature is too low, the material remains porous. The areas with the loosely packed grains crack open further during the sintering process and often remain as structural faults and therefore as potential defective areas in the framework. After sintering, a slow cooling rate is essential to minimize nonuniform residual stress. Slow cooling is particularly important when a structure has asymmetric distribution of thick and thin structures that cool at uneven rates. Therefore, zirconia is highly sensitive to preparation methods, handling conditions, moisture, heating profile, and cooling rate. If any of these are uncontrolled such that the bulk of the material is transformed into the monoclinic phase, the mechanical properties of zirconia become compromised, increasing the risk of spontaneous catastrophic failure.

It must be re-emphasized that in general the main disadvantage of ceramics as compared to metals and polymers is their low fracture toughness, which measures a material’s resistance to crack propagation under applied stresses. In ceramics, and specifically in zirconia ceramics, finer grain crystalline structure results in increased toughness and hardness as well as better wear resistance. Grain size also determines the surface finish quality achieved by grinding and polishing operations. Fine grain structure allows the size of the surface microasperities to be decreased after the surface finish operation, resulting in a lower coefficient of friction and a smoother surface that will be less prone to retain foreign particles. Wear occurs by many mechanisms that are beyond the scope of this chapter, and one of the common wear mechanisms involves the combination of tribocorrosion and microcrack formation, propagation, and microfatigue fracture.

Currently, a number of commercially available zirconia dental implants are available in the market in a range of widths and sizes for various edentulous sites. Manufactures also supply removable overdenture attachments based on the classic ball-type attachments (Z-Systems) and Locator attachments (Zeramex); however, these zirconia-based attachments have not yet been investigated independently with published in vitro or in vivo studies. Zirconia-based dental implants are still at an early stage and lack support for their long-term use. Nevertheless, in vitro studies are promising and support the possibility of their use as an alternative to titanium implants. The three factors that need to be satisfied so that zirconia dental implants may be considered as an alternative to titanium implants are (1) osseointegration and its long-term maintenance, (2) prosthodontic restorability, and (3) fatigue and fracture resistance. These factors are discussed in the subsequent sections.

Osseointegration and its maintenance

Osseointegration to a titanium surface is well documented and is the gold standard for all other materials. Biologically, zirconia is an excellent material for osseointegration and soft tissue health. In vitro studies have confirmed that zirconia does not provoke inflammatory, allergic, immunologic, or carcinogenic reactions in various cell cultures.16 Zirconia is known to be an osteoconductive material. Animal studies comparing zirconia implants to titanium implants using histomorphometric evaluation of bone-to-implant contact (BIC) report no statistically significant difference between both groups at any healing time point.17–20 Furthermore, studies also show that zirconia and titanium have improved BIC and removal torques with modified roughened surfaces compared to smooth machined surfaces.18,21,22

Clinical evidence from sound long-term scientific clinical trials for the biologic compatibility and osseointegration of zirconia implants is lacking. Much of the literature reports case reports and case series,15,23–29 with a very small number of short-term clinical trials25,28,30–40; however, none are randomized controlled trials looking at zirconia and titanium implants, and most studies only provided evidence on a low level due to study design limitations.

At this stage, it is not possible to make conclusions about whether zirconia implants are a suitable substitute to titanium implants or if they will behave in the same way as alumina- based implants. However, it is worth noting that fracture of zirconia dental implants is not an uncommon occurrence and has been reported to occur during surgical placement39,41 and after prosthetic loading.32,36,38,42 To date, reported clinical survival rates of zirconia implants range from 82.4% to 100% over 1 to 5 years of follow-up, but when strict success criteria that take into consideration radiographic marginal bone loss are applied, success rates decrease significantly. For instance, Kohal et al34 investigated one-piece zirconia implants and found that 48% of implants restored with single crowns and 51% of implants restored with fixed partial dentures had lost 2 mm or more marginal bone at the 1-year follow-up. The authors therefore concluded that that the one-piece zirconia implant design could not be recommended for clinical use because success rates were significantly lower than conventional two-piece titanium implants.43 Nevertheless, there is speculation that aspects of implant design or temporary cement entrapment contributed to the high marginal bone loss rather than a direct consequence of the material being zirconia.

Prosthodontic restorability

One-piece zirconia dental implants lack abutment flexibility and screw retention; therefore, restorations must be cemented (Fig 4-1). This means there is no option for retrievability of prosthetic components and no flexibility in choosing from a variety of abutment types and angles. This can result in problems with esthetics and emergence profiles when implants are unfavorably placed in the esthetic zone. Improving implant gingival esthetics by developing soft tissue contours and papilla through tissue modeling using a series of provisional restorations is also not possible with one-piece implants.

Fig 4-1 (a to d) A one-piece zirconia implant system. These designs require the implant to be perfectly positioned. The prostheses are not easily retrievable because they are cement retained.

Adjustment of the abutment portion of an implant immediately after surgical placement as recommended by manufacturers to provide occlusal clearance is possible but in some instances can contribute to problems with one-piece implant systems. Grinding with rough diamond burs introduces deep surface flaws and creates crack initiation sites that may decrease the clinical life expectancy of the implants, and vibration trauma from grinding could adversely affect primary stability of an implant, particularly because zirconia is extremely tough and very difficult to section or shape. Furthermore, one-piece implants following placement are directly or indirectly immediately loaded, whether they are restored or not. Forces from masticatory activity and tongue pressure onto the abutment portion of one-piece implants are inevitable and unavoidable during the initial healing phase. As a result, case selection must be highly selective when one-piece implants are to be used. This is based on the findings from the limited clinical trials available that show a high failure rate for one-piece zirconia implants when placed in immediate extraction sites as well as grafted sites.25,32

In the meantime, manufacturers suggest immediate provisionalization of one-piece zirconia implants when the final insertion torques exceed 30 to 35 Ncm, but when implants are placed with an insertion torque lower than 30 Ncm, it is recommended that the implant be splinted to the neighboring teeth to reduce the risk of micromovement of the implant during the initial healing phase. This method of primary stabilization of one-piece zirconia implants has been found to increase the risk of early failure by some clinicians,25 but not others.26,31 It might be speculated then that the periodontal health of the adjacent teeth and their mobility might have a direct influence on the success rate of these splinted one-piece zirconia implants. This poses a serious limitation for these systems and again highlights the importance of meticulous patient selection. Because of this, some manufacturers are now marketing two-piece zirconia implants. However, the abutments are bonded in place either before or after surgery using a permanent bonding resin-based system, resulting in a lack of prosthetic flexibility for interchangeable abutments.27,39,40 As a result, these two-piece zirconia implant systems have just as limited clinical retrievability as one-piece implants and do not offer the prosthetic flexibility of two-piece titanium implants.

Fatigue and fracture resistance

Among dental ceramics, zirconia is perhaps the most suitable for use as a substitute for titanium dental implants because of its mechanical properties. Zirconia has very high flexural strength (900–1,200 MPa), but its elastic modulus is twice that of titanium. In vitro studies suggest that zirconia implants are able to withstand maximum loads in the oral cavity, even after artificial aging.44–46 Zirconia has a high elastic modulus but is extremely brittle compared with metals that are highly ductile. This explains why zirconia cannot withstand much strain before fracturing compared with titanium. Coupled with low fracture toughness, the main concerns with zirconia implants are their susceptibility to catastrophic implant fracture over time, similar to that observed with alumina-based implants, and their higher likelihood for stress shielding.

CAD/CAM technology enables the reproduction of various ceramic geometric shapes with excellent marginal fit.47,48 However, the brittle nature of zirconia precludes it from being used in thin sections under loads, such as high clamping. High stress concentrations in this region place the zirconia joint system at risk of catastrophic fracture.49,50 Machining of screw threads within zirconia implants for a prosthetic screw poses a potential risk of fracture of zirconia implants at the level of the screw threads. For this reason, zirconia implant manufacturers are mostly limited to the production of one-piece zirconia implants to eliminate thin fragile features within the implant and to reduce the risk of fracture.

Zirconia is also highly sensitive to moisture and low- temperature degradation (LTD), although to date there is no accepted mechanism to explain LTD.51 How great a problem this is for zirconia implants is still unknown. The vulnerability of zirconia to aging is compounded by the fact that the severity differs among different types of zirconia and even in zirconia from the same manufacturer that has been processed differently.52–54

Surface treatments such as particle air abrasion,55 polishing,56 and multiple firing cycles57 have been found to directly influence the strength of zirconia material both positively and negatively, demonstrating its unstable and sensitive nature. During the CAD/CAM milling procedure, zirconia is subjected to surface damage, which can significantly reduce its overall flexural strength and service life.55 Air abrasion with aluminum oxide (Al₂O₃) of different grit sizes has different effects on the flexural strength of zirconia depending on the Al₂O₃ particle size. Zhang et al found a reduction in the flexural strength of polished zirconia after air abrasion with 50-µm Al₂O₃,56 compared with Wang et al55 who found flexural strength was increased following air abrasion with 50-µm Al₂O₃. It has been suggested that air abrasion increases the flexural strength of zirconia by creating a residual compressive stress field on the surface of the implant. This occurs as a result of a tetragonal-to-monoclinic transformation, by removal of weakly attached surface grains, and by eliminating the trace lines produced during the CAD/CAM milling process, thereby reducing surface roughness.58 However, this process is sensitive to the particle size and pressure of the air abrasion procedure. Severe surface damage in the form of sharp scratches, cracks, and grain pullout causes an increase in the surface roughness of zirconia, and a decrease in the flexural strength was observed when 120-µm particles were used.55 In the meantime, polished zirconia surfaces have higher flexural strengths than nonpolished or rough surfaces.55,56 This is due to the reduction in severity and number of surface defects and flaws by the polishing process59 and illustrates the strong correlation between flexural strength and severity of surface damage.

As mentioned previously, microroughened zirconia implants have shown improved BIC and better osseointegration compared with machined zirconia implants. Surface modification of machined zirconia implants has been suggested by using air abrasion,21 acid etching with a hot solution of hydrofluoric acid,25,60 surface coating with bioactive ceramics,25,61 sintering with a slurry containing zirconia powder and a proprietary pore former,62 selective infiltration etching to produce nanoporous surfaces,63 and femtosecond laser microtexturing modification techniques.64,65 While these roughened surfaces all seemed to improve integration in bone, it is not yet known if they impose any adverse effects on the mechanical properties of the zirconia implants, and at this stage, there is no consensus on the best surface treatment method for zirconia implants.

Zirconia dental implants have also been reported to fracture during surgical placement and after prosthetic loading.32,36,38,39,41,42 In one particular study, 3.25-mm-diameter zirconia implants suffered significantly higher fractures rates compared to implants with 4 mm diameter or larger, and 92% of fractured implants occurred in the anterior regions.36 Fractographic analyses revealed notches and scratches on the surface of the zirconia from the sandblasting roughening process at the origins of fractures. All fractured implants had similar crack propagation patterns, originating from the palatal aspect, with the cracks heading toward the buccal surfaces. These fractures were explained to be due to singular bending loads after prosthetic loading.42 These fractures were concentrated in the anterior regions, which may indicate that zirconia dental implants are potentially more susceptible to off-axis loading complications than titanium implants. This should be of no surprise, because ceramics are brittle and respond catastrophically to high bending forces compared to metal alloys. While a narrower implant diameter may also be a contributing factor, the fracture rate of nearly 10% after a follow-up period of 36.75 months was higher than the low fracture rates reported in the literature for titanium implants.66–68 An interesting observation with most fractured titanium implants is that fracture is time dependent, with the incidence of fracture increasing the longer the loading period. Fracture analyses showed that metal fatigue was the main process of failure, and a number of causes have been suggested, such as peri-implant bone loss, unfavorable implant orientation, overloading, and ill-fitting prosthetic frameworks. These observations have been made after years of follow-up of titanium implants, and only time will show how these factors will contribute to the incidence of zirconia implant fracture.

So far, zirconia has been shown to be at least equivalent to titanium as an implant material in terms of its biologic compatibly, with excellent hard and soft tissue responses. However, mechanical properties such as fracture resistance, fatigue life, reparability, compatibility with common dental adhesives, limited prosthetic options, and even higher modulus mismatch make results less favorable with zirconia implants, and they are yet to be clinically confirmed to be able to predictably withstand long-term oral function.

Summary of implant materials

Titanium materials have long been considered to be bio- inert, and it has been believed that the presence of a titanium implant in an osteotomy site should not influence the wound healing process. While the mechanistic elucidation is not complete, it is increasingly clear that osseointegration is not achieved only via bone formation. Recent observations and experimental evaluations indicate that there are distinct molecular and cellular behaviors that appear to be unique to peri-implant tissue. Some of these characteristics contribute to the mechanical advantage and long-term stability of osseointegrated implants. In addition, peri-implant bone may not undergo the same biologic and pathologic sequences as tooth-bearing alveolar bone. The maintenance of osseointegration may require special consideration. This is in part because similar results for marginal bone loss were reported for NobelDirect one-piece titanium implants (Nobel Biocare) by a number of authors.69–71

In terms of soft tissue response, zirconia may have an advantage over titanium in its reduced affinity to plaque formation.72–75 Indeed, healthy soft tissues have been consistently reported in case reports and clinical trials using zirconia implants.24–28,32–34 However, long-term survival of zirconia implants is needed before they can be recommended by the authors. The longest published data on zirconia implants reported a 95% success rate of 831 one-piece CeraRoot zirconia implants placed in a private practice setting after 5 years of follow-up.25 It is important to note that radiographic bone loss and soft tissue parameters were not investigated in this study, and therefore only survival rates could be established.

Prosthetic Materials

Full-arch suprastructure materials

A discussion of full-arch suprastructure materials must be preceded by a review of the mechanical properties that best enable the practitioner to select the most appropriate material for the treatment plan design. For this purpose, the 0.2% proof stress/yield stress and modulus of elasticity are two mechanical properties that best inform the practitioner of the elastic stress range within which the material will be able to flex and return to its original dimension and the degree of stiffness of the material compared to other materials or the supporting oral structures that surround the implants onto which the suprastructure is attached. The 0.2% proof stress/yield stress is the stress value (MPa) below which the material is elastic and above which the material becomes plastic (will distort) or in the case of high elastic modulus/brittle materials such as zirconia will fail. The elastic modulus is the ratio of tensile stress to tensile strain on the linear portion of the stress/strain curve and denotes the stiffness of the material.

When selecting an appropriate alloy or ceramic such as zirconia for long-span suprastructures that are going to be subject to heavy loads, the clinician should consider a material with a high elastic range (0.2% proof stress/yield stress) combined with a low stiffness (elastic modulus). The higher the elastic modulus of the suprastructure, the greater the occlusal forces that will be transferred through to the supporting implant fixtures and the osseointegrated implant–bone interface.

Type IV alloys

Silver palladium, gold palladium, and gold platinum palladium alloys were popular in the past for the fabrication of long-span implant suprastructures. Silver palladium alloys have the advantage of having a low density (typical density of 11.4 cm3) compared to gold palladium alloys (typical density of 14.5 cm3) and gold platinum palladium–based alloys (typical density of 18.1 cm3). This means that the amount of silver palladium alloy required to cast a suprastructure pattern is 21% less than gold palladium alloy and 37% less than gold platinum palladium alloy. This translates to considerable savings for the overall cost of the casting because less alloy in terms of price per gram is required. In terms of 0.2% proof stress, gold palladium alloys have the higher proof stress (572 MPa) compared to silver palladium (531 MPa), followed by gold platinum palladium (448 MPa). The main disadvantage of these noble alloy systems is their high price per gram in the current market.

Cobalt-chromium and type V titanium alloys offer significantly cheaper options. Type V titanium has a higher proof stress of 862 MPa compared to cobalt-chromium of 690 MPa, both of which are superior to the noble alloys. While cobalt-chromium can be milled and cast using typical laboratory techniques, type V titanium must be milled. The lower modulus of elasticity of type V titanium (120 GPa) compared to cobalt-chromium (220 GPa) makes it easier to mill. In summary, type V titanium has the most advantageous mechanical properties combined with low cost per gram. The only disadvantage is the requirement to mill the framework via a CAD/CAM system.

Monolithic zirconia

Prostheses made of 3Y-TZP ceramics are exposed to moisture and continuous dynamic loading. These conditions essentially result in a decrease in the values for measured strength, resulting in hydrothermal aging of the material. This decrease in strength can be simulated in an autoclave at 130 to 150°C and a water vapor pressure of 10 bars.76 The aging process in vivo takes place far more slowly, as dental restorations are never exposed to such extreme conditions. The addition of approximately 0.5% Al₂O₃ to ZrO₂ substantially increases the resistance to hydrothermal aging. Degradation and pitting of the surface as a result of hydrothermal aging appear to be highly unlikely, but this needs to be assessed in long-term clinical studies.77

Despite these limitations, for the past decade, zirconia has been used routinely in copings and frameworks to support feldspathic porcelain. Inconsistencies in clinical reporting of outcomes have further confounded understanding of the appropriate parameters for reliable fabrication of dental restorations with this material, such as framework design and proper support of the laminated porcelain and cooling rate of the laminated porcelain-zirconia substructure. Long-term clinical trials have confirmed the success of functionally veneered zirconia in the anterior zone78 and selected posterior applications.79,80 However, chipping and fracture of the feld-spathic porcelain veneer continue to be issues the clinician must consider when using this material, as with other bilayered restorations traditionally used in dentistry.

Porcelain delamination had not been reported to be a problem until sintered zirconia started being widely used in the United States, even though Europeans had been fabricating porcelain-fused-to-zirconia (PFZ) restorations on zirconia manufactured with HIP since 1993. It was found that slow heating and slow cooling was the key to eliminating porcelain delamination.

More recently, the use of monolithic zirconia prostheses with no feldspathic porcelain veneer in areas of high occlusal stresses and minimal feldspathic porcelain veneer on the facial surface of zirconia to enhance esthetics, or replace gingival structures with pink feldspathic porcelain, is becoming popular. This is due to the development of efficient milling procedures and the development of zirconia blocks with enhanced esthetic properties (see chapter 8, section entitled “Monolithic Zirconia Prostheses”; and chapter 10, section entitled “Monolithic zirconia prostheses”). Follow-up times are relatively limited, but the initial data look promising.81–84

Dental zirconia enjoys properties (Box 4-1) that enable this material to serve as a substitute for traditional restorative materials. There is a greater demand for esthetic and reliable materials for our patients, and the availability of zirconia may prove to be an appropriate choice.

Box 4-1 Properties of dental zirconia

• High strength and toughness85

• Favorable wear resistance86

• Minimally abrasive87

• Acid resistant

• Thermal insulator

• Highly biocompatible16

• Low plaque adhesion74

• Flexible design possibilities for all applications

• Esthetic88

• Maintains occlusal stability86

Framework creation. Zirconia frameworks must be properly designed with strict adherence to design protocols (see chapter 10, section entitled, “Procedures for zirconia frameworks combined with individual lithium disilicate crowns”). The size of the connectors of the prosthesis should be maximized and should be thick and broad. A minimum area of 9 mm2 should be available for connectors, and framework wall thickness should be no less than 0.8 mm (Fig 4-2). The frameworks must be anatomically shaped to ensure even layer thicknesses of any veneering ceramic. Insufficient support for the veneering ceramic on the coping or bridge framework can trigger chipping and fractures. Sintered zirconia is very dense and highly acid resistant, which prevents the absorption and surface accumulation of pollutants. As a result, zirconia fixed partial denture pontics in contact with the tissue do not need to be veneered with porcelain.

Fig 4-2 (a and b) Zirconia frameworks must be designed with strict adherence to design protocols. The connectors should be thick and broad. (Courtesy of Dr A. Pozzi.)

Full-contour monolithic zirconia restorations are becoming popular due to the availability of more translucent formulas. Also, improved coloring (presintering) and glazing (postsintering) techniques have been developed. Monolithic restorations (Fig 4-3), as opposed to laminated porcelain restorations, are preferred on functional surfaces, because the porcelain veneering is predisposed to chipping and fracture. The primary objection to the use of monolithic zirconia restorations is based on the perception that the functioning surfaces become abrasive following adjustment, leading to excessive wear of the opposing natural dentition. However, if the functional surfaces are properly polished with the appropriate finishing instrumentation, they can be effectively smoothed, eliminating the risk of excessive wear of the opposing tooth surfaces. The result is a surface that is very dense, smooth, stable, and minimally abrasive. Excessive heat generation during adjustment and/or polishing should be avoided to minimize the risk of changes in the crystalline structure of the surface, potentially predisposing the material to chipping and fracture.

Fig 4-3 (a and b) A full-arch zirconia prosthesis. The functional surfaces are monolithic, whereas the facial surfaces in the esthetic zone are restored with laminated porcelain.

Zirconia dental prostheses can be colored in the presintered stage with special coloring liquids (stains). The result is a sintered restoration in basic dentine color. Recently, precolored zirconia blocks have become available, and coloring zirconia in the presintered stage does not affect the strength of the material. This technique can also be used with newer translucent blocks to customize the surface of monolithic restorations.

In porcelain-zirconia restorations, the feldspathic ceramic (porcelain) must be sufficiently supported, especially because there is not a chemical bond between the feldspathic ceramic (or any other ceramic or composite) and the zirconia. The process of applying porcelain onto zirconia substrates is different than that used for traditional metal-ceramic restorations. The laminated porcelain is retained by compression and micromechanical adhesion to zirconia. Sintered zirconia is an extremely poor heat conductor; therefore, different heating and cooling rates should be used when applying ceramic onto zirconia and/or during the glaze cycle. This is particularly important when fabricating large-volume structures, and these calculations are based on the weight of the structure.

Processing. Most zirconia restorations are created from the green (partially sintered) state. During sintering, the prosthesis framework is suspended upright in order to control volumetric shrinkage of the material (see Fig 4-3). In this stage, zirconia can be easily milled and ground. However, once sintered, zirconia requires careful handling due to its poor heat conductivity. Densely sintered zirconia should not be subjected to extensive dry grinding; otherwise, there is a danger of overheating the material, which creates surface tensions through distortion of the crystal lattice, fissures, and later, fissures in the facial surfaces of the prosthesis. Densely sintered zirconia should therefore only be altered in a moist environment. The use of coarse diamonds, excessively high rotational speeds, and excessive pressure must also be avoided. Sintering ovens cannot “restructure” the zirconia or “heal” the cracks.

Although transformation strengthening compresses the cracks that have been produced, this so-called “self-repairing mechanism” only works once. If another crack is produced in the same region, it can no longer be compressed. In addition, the monoclinic crystalline phase that results from the transformation has a higher coefficient of thermal expansion (CTE) than the tetragonal crystalline phase. The outcome is that the CTE of the framework may be altered and may not be compatible with the CTE of the facing (laminated) ceramic; hence, the survivability of the definitive restoration may be compromised.

For one-piece, full-arch zirconia prostheses comprised of suprastructure and teeth, veneering porcelain is sometimes added for esthetic results. However, chipping and delamination of veneering porcelain and ceramic core fracture are persistent issues with all-ceramic restorations,89–92 and this can have serious implications if it occurs with the all- ceramic implant restorations. A meta-analysis of single- implant crowns reported significantly higher 5-year survival rates for metal-ceramic crowns than all-ceramic crowns.93 It was also found that chipping of the veneering porcelain in metal-ceramic restorations was significantly higher with crowns restoring implants than with crowns restoring the natural dentition due to the lack of a periodontal ligament (and proprioception) with implants.94 This suggests caution when considering using all-ceramic restorations on implants that do not offer simple prosthetic retrievability.

Frequent causes for fractures of zirconia frameworks include the following:

• Inadequate framework wall thickness (min. 0.8 mm)95

• Connectors too small (9 mm2 three restorative units [span], 12 mm2 four restorative units [span])

• Thermal shock introduced by dry grinding the densely sintered framework, use of coarse diamonds, excessive rotational speeds and/or exerting excessive pressure58

• Thermal shock introduced by inappropriate cooling and heating rates57

• Nonpassively fitting zirconia restoration/framework

Fracture and delamination of ceramic facings may be secondary to the following:

• Insufficient support for the laminated ceramic by the coping or bridge framework96

• Copings or frameworks that are not anatomically shaped96

• Uneven thickness of laminated porcelain

• Inappropriate porcelain firing procedure

• Use of conventional ceramic heating and cooling rates

• Use of inappropriate equipment not suited to the material

Restorative workflow. The procedures used to create zirconia dental restorations do not use traditional methods (see chapter 8, section entitled “Monolithic Zirconia Prostheses”; and chapter 10, section entitled “Monolithic zirconia prostheses”). The fabrication of zirconia restorations requires dental technicians to master new and rigorous skill sets to create reliable and esthetic outcomes. There are currently two types of zirconia blocks available for dental applications: (1) fully sintered zirconia blocks (HIP zirconia blocks) and (2) partially sintered zirconia blocks (CIP zirconia blocks). The fully sintered block is generally stronger due to the increased alumina content. Unfortunately, the alumina imparts a white, chalky appearance and is not particularly suited to esthetic applications. This type of zirconia is primarily used to fabricate dental implants and implant abutments as opposed to definitive esthetic restorations.

Milling fully sintered zirconia is technique and time intensive. Color changes are not possible within the body of the zirconia; therefore, surface treatments are required to create an acceptable esthetic outcome. This historically has been accomplished with the addition of feldspathic ceramic. For most applications, other than anterior implant restorations, this approach has not resulted in clinically reliable outcomes because of the undesirable incidence of chipping and fracture of the laminated porcelain.

Partially sintered zirconia blocks can be milled efficiently and are amenable to shade correction and characterization before or after full sintering. There is a minor decrease in strength with this material, but it is not clinically significant. The major advantage is that the potential for shade correction and characterization eliminates the need for feldspathic ceramic on functional surfaces. Pink feldspathic ceramic is also routinely used to substitute for lost hard and soft tissue. This material is not in an area of function, so it is not subject to chipping and fracture.

Polymethyl methacrylate and nanohybrid composites

The typical flexural strength of polymethyl methacrylate (PMMA) ranges from 66 to 131 MPa,97 which limits its use as a suprastructure material unless it is well supported by a metallic framework such as in a fixed hybrid prosthesis. It can be used as a temporary material if confined to the anterior region (Fig 4-4). Likewise, the use of nanohybrid composite materials is limited due to their relatively low flexural strength that ranges from 76 to 147 MPa. These flexural strength values have been reported to have decreased by as much as 40% after 30 days exposure to artificial saliva.98

Fig 4-4 (a to c) Milled PMMA prototype prostheses. They are designed to be used for 2 to 3 months and will eventually be scanned in preparation for fabrication of monolithic zirconia prostheses.

Polyaryletherketone (PAEK) polymers such as polyetheretherketone (PEEK) and polyetherketoneketone (PEKK) have much better mechanical, chemical, water resistance, and biocompatibility properties than PMMA. Both thermoplastic materials can be thermoformed and readily machinable with CAD/CAM equipment and dental handpieces, and the ease of processing along with desirable properties is gaining increasing applications in dentistry and medicine. Both have elastic modulus that is much closer to bone. Although PMMA offers better optical properties and esthetics, PEEK and PEKK can both be processed into acceptable tooth colors (albeit with more opaqueness than PMMA) and can more easily be modified and repaired than zirconia (Table 4-2). Both are polymers of the PAEK family. The wider melting range of PEKK makes it more practical for 3D printing and melt processing without sacrificing mechanical properties.

PEEK (-C₆H₄-OC₆H₄-O-C₆H₄-CO-)n is a crystalline polymer chain comprised of an aromatic backbone with interconnected rigid ketone and flexible ether groups. Because it has no metallic elements that can leach out of the matrix to induce allergies or cytotoxic or genotoxic activities, it has been used in FDA-cleared implantable medical devices. Because PEEK is stable up to 335.8°C, PEEK implants are stable under various medical sterilization processes such as gamma irradiation, steam autoclave, and ethylene oxide.

By replacing the flexible ether group with a rigid ketone, PEKK has higher glass transition temperature and mechanical properties. Furthermore, the second ketone group can be either straight or angled with respect to the first ketone group, and the mixing ratio of straight PEKK versus angled PEKK can allow further fine tuning of thermal and mechanical properties. Structurally, the mixing of straight and angled segments reduces crystallinity, depending on composition and cooling rate, to produce a combination of amorphous and crystalline phases. Besides increasing flexibility in the thermal processing window, PEKK also improves the mechanical properties and dimensional stability over PEEK due to reduced crystallization shrinkage secondary to having fewer crystals. Amorphous PEKK is also somewhat less opaque than PEKK. Both have been reinforced with nanofibers and nanoparticles to further improve mechanical properties and have been surface modified to enhance interactions with biologic cells and tissues. Based on the rapid progress in polymer engineering, more versions of nanoengineered PEEK and PEKK will be available for dentistry soon.

Substructure materials

Titanium

As well as being commonly used for fabrication of implant fixtures, the properties of titanium also make it suitable for use in fabrication of prosthetic components. Titanium is relatively low cost and has low density, yet possesses high mechanical strength and an oxide layer that is highly protective against the oral environment.99 Despite these desirable features, the use of titanium for particular prosthetic solutions introduces new issues, one of which is difficulty in handling when investing and casting into the required prosthesis. It requires a high melting temperature and is prone to incorporating porosities during the casting process.99,100 Conventionally cast titanium has also been found to have inferior accuracy due to shrinkage.101 Accuracy and strength are both highly important when fabricating implant prostheses, particularly with multiunit prostheses. Hence, use of conventionally cast titanium as a base material is not a practical or predictable option. The introduction of CAD/CAM has made the fabrication process much more straightforward, contributing to the increasing use of titanium for the manufacture of implant prostheses.102 Several companies now offer milling services for the fabrication of implant prostheses using various materials, including titanium.

Titanium is not suitable for use as a monolithic implant prosthesis due to its metallic appearance. Therefore, it can only be used as an implant prosthesis substructure that is layered with an esthetic veneering material such as porcelain. One option is for direct layering of material onto a titanium substructure, and this option becomes a more attractive option when cost is an issue. The weakest part of the prosthesis then becomes the bonding interface, and the overall success of the prosthesis becomes much more reliant on the bonding characteristics of the veneering material onto titanium. When using acrylic bonded to titanium as a veneering material, the prosthesis performs satisfactorily in in vitro studies, particularly with the addition of material-specific metal primers.103 This combination of acrylic and titanium also performs well when used clinically.104

A more esthetic option would be to layer the titanium substructure with porcelain. However, this option must be used with caution, as the thick oxide layer on titanium creates issues with the bonding to porcelain, mainly when the porcelain is conventionally layered and fired onto the metal surface as an additional firing process further increases the thickness of the metal oxide layer. The thick oxide layer leads to an oxide-oxide failure being observed when the porcelain-fused-to-titanium prosthesis is subjected to mechanical forces, which may indicate a weak bond strength. Peer-reviewed studies also back those observations, showing that there is minimal porcelain bonding to the metal when examining the failure surface under scanning electron microscopy.105

Type IV noble alloys

Type IV noble alloys still offer the best solution to implant substructures, especially when bonding to porcelain is required (see chapter 8 section entitled “Metal-Ceramic Fixed Prostheses” and chapter 10, section entitled, “Procedures for metal-ceramic fixed prostheses”). They have a superior metal-ceramic bond strength due to the relative lower modulus of elasticity of the noble alloys compared to the base metal alloys, such as nickel-chromium and cobalt-chromium, resulting in less residual stress at the porcelain-metal interface from the compression bonding component of the metal-ceramic bond. In addition, the chemical bond component is superior in that the noble alloys contain only sufficient base metal elements to facilitate optimum chemical bonding. However, the propensity of the base metal alloy to oxidize readily often results in oxide-oxide failure at the bond interface, especially if there have been multiple firings of the porcelain veneer. In addition, the color of the base metal elements included in the noble bonding alloys to provide chemical bonding offer the ceramist a more suitable oxide color that more easily facilitates the achievement of good esthetic outcomes. The only negative aspect of type IV noble alloys is their high cost per gram compared to base metal alloys, which also have a lower density than noble alloys.

Overdenture bar materials

Titanium

Traditionally, to ensure appropriate fit when casting an overdenture implant bar, several prosthetic components were required. The use of accurate CAM methods has reduced the cost and complexity for the fabrication process of an overdenture implant bar. Titanium, with its relatively low cost and biocompatibility, is an attractive metal to use for that purpose,106 but only when the implant bar is designed to be implant-supported (Fig 4-5; see chapter 9, section entitled “Implant-Supported Overdentures [Fixed Detachable Prostheses]”). Titanium is not suitable for implant-assisted bar designs because of its propensity for wear (see chapter 9, section entitled “Implant connecting bar—CAD/CAM techniques”).

Fig 4-5 An overdenture implant connecting bar designed to be implant- supported and milled of titanium. This bar was milled to a 3-degree taper, and retention was supplemented with low-profile individual attachments (see chapter 9).

The most common issue faced when using implant overdentures is the loss of retention from the attachments. Although at times it is merely due to the frequent insertion and removal of the overdenture for cleaning and maintenance, the rate of retention loss may also be exacerbated by improper design of the attachment system and lack of adherence to basic denture design principles (see chapter 7). Hence, the design of the overdenture should follow basic design principles for a conventional denture, especially when implant-assisted designs are used, and the chosen attachment mechanism should have enough fatigue retention to avoid the frequent need for replacement by the clinician. When choosing an attachment system for the overlaying acrylic denture matrix onto the implant overdenture, there are several options available (see chapter 7).

Cobalt-chromium alloys

Cobalt-chromium alloys can be both cast and milled via a CAD/CAM system (Fig 4-6). The milling process has the advantage of being better able to achieve passive fit than the casting process (see chapter 7, section entitled “Design and fabrication of the implant connecting bar: CAD/CAM method”). This material is favored when fabricating implant-assisted bar designs because of its resistance to wear.

Fig 4-6 (a and b) A digitally designed implant- assisted overdenture implant connecting bar milled from cobalt-chromium.

The wax pattern and sprue system can be prone to contraction and distortion prior to investing and the subsequent shrinkage stresses involved with a long-span structure on cooling. The framework often has to be cast in sections and subsequently laser welded together to achieve passive fit.

Type IV noble alloys

Overdenture bar systems can be cast, but with the same issues as outlined for titanium and cobalt-chromium alloys. The adaptation of preformed type IV gold bars—either U-shaped or ovoid—to preformed nonoxidizing gold cylinder abutments is not recommended. This technique offers a fairly simple technique that the dental laboratory can follow to produce a passively fitting overdenture bar system without requiring sophisticated laser welding or milling machines. However, evidence from an evaluation of simple laboratory failure modes for distal cantilevers has shown (1) that in soldered joints used for overdenture bars, cracking was initiated in the solder due to fatigue, regardless of the type of soldering material used, and (2) that the joints have relatively low yield stresses and were prone to plastic deformation under maximum occlusal forces.107 This study also reported that corrosion fatigue, in conjunction with masticatory cyclic loading, appeared to be the primary factor in the etiology of observed clinical failures. A study by Waddell et al108 analyzed clinical failures of soldered bar attachment systems and showed that corrosion, followed by corrosion fatigue, appeared to be a key factor in the onset of the failure process.

Computer-aided manufacturing

CAM involves a computer-controlled manufacturing machine following instructions from a CAD program for the fabrication of structures. The increased use of CAM technology is due to ease of workflow, improved accuracy of scanning technology, higher demands to manipulate material, and necessity for more accurate custom parts for various applications, including surgical and implant dentistry.109 There are many reported advantages of this technology when compared to traditional casting and manufacturing methods, including the reduction in manual handling and labor-intensive steps in the fabrication of prostheses using a predictable workflow. Other advantages include reduced production time (hours versus days), lower overall cost (in terms of labor), simpler quality control, less reliance on manual skills of the technician, and more control over the microstructure and properties of the final product.110 While CAM has been used as a synonym for subtractive manufacturing, it does not specify the method used.111 Two main CAM techniques are used in dentistry: subtractive (milling) and additive (3D printing).

Milled materials

Subtractive manufacturing utilizes computer-controlled movements of cutting tools to reduce a single block of material into the designed structure. Subtractive manufacturing machines are distinguishable through the number of milling axes they use, mostly 3, 3+2, and 5 axes of movements available. While more axes do not necessarily guarantee an increase in surface finish, more axes usually allow faster milling, less operator dependence, more complex tooling path, more complex shapes, and more accuracy due to elimination of indexing and reorientation.111 In addition to the number of axes a milling machine has, another variability is the type of machining involved; this is usually categorized as hard and soft machining.112

Hard machining involves the milling of the highly sintered form of the base material. This type of machining tends to subject the material and the milling components to a very high amount of force, increasing temperature and wear on the milling burs. The effect is particularly evident when milling hard base materials such as cobalt-chromium and zirconia. In addition, there is a particular problem with low conductivity materials, such as zirconia, in which quick temperature variations may cause surface microcracking and plastic deformation on a microscopic scale.113 While that does not change the accuracy of the fit of the prosthesis, the residual stress does increase the risk of low-temperature degradation and may reduce the lifespan of the prosthesis.114 This process produces a very dimensionally accurate prosthesis that requires significantly less post-processing115; however, bur wear not only increases production cost and time due to bur failure, but the milled structure can also suffer from microcracks that are not healed because there are no subsequent sintering steps.

Soft machining, on the other hand, involves milling presintered blocks of material to an enlarged net shape and then sintering to the final dimensions. Unlike hard machining, the softer nature of the base material reduces the overall stress on the milling components and hence reduces the need for cooling and increases the lifespan of the cutting tips.115 The structure is then subjected to a sintering step to achieve the properties required for use, and some microcracks can be eliminated during the sintering process. The sintering step causes shrinkage in the milled structure, which can reach around 25% for zirconia, and the manufacturer’s specified shrinkage has to be accounted for by the CAD software prior to milling.114 However, the risk of damage to surface and wear has to be weighed against the accuracy sought from the final structure.116 This was highlighted in another study comparing zirconia frameworks made from fully sintered and presintered blocks.117 Kohorst et al found that the milling of sintered zirconia to fabricate restorations produced a prosthesis with a significantly better fit than prostheses made from milling then sintering presintered zirconia blocks.117 The types of materials that are commonly used in subtractive manufacturing include ceramics, resins, and metals (Fig 4-7).

Fig 4-7 Scanning electron micrographs of hard-milled (a) and soft-milled (b) materials at 1000× magnification. Note the high porosity on the soft-milled specimen.

PMMA is a highly dense acrylic-based material commonly used to fabricate provisional prostheses for use in the oral cavity. However, conventionally mixed PMMA is prone to shrinkage and produces heat while polymerizing. The reaction is uncomfortable for the patient, and the final product is not as predictable in regard to being defect and porosity free. When milled, PMMA has relatively high fracture resistance and chemical stability,118 making it an ideal material to use for provisional restorations when extensive prosthodontic treatment is being undertaken (see chapter 8, section entitled “Monolithic Zirconia Prostheses” and chapter 10, section entitled “Monolithic zirconia prostheses”). In implant pros- thodontics, PMMA is also used as a prototype for a planned prosthesis, allowing for examination and adjustment of esthetics and occlusion of the final prosthesis before fabrication of the definitive prosthesis, potentially saving time and effort for both clinician and technician.

Milling of zirconia and lithium disilicate crowns using CAM technology is a highly popular procedure. The use of titanium bases specific to the implant interface allows for milling of direct-to-fixture–type ceramic restorations. Both types of ceramic have been found to have suitable strength for use as implant crowns.119 Special attention has been given to the properties of zirconia and its polymorphic characteristics in preventing crack propagation. Its use has increased with the spread of computer-aided milling due to the relative ease with which milling machines produce a structure made from zirconia compared to traditional production methods.112 Zirconia implant abutments have a high success rate in the anterior region, although they are not recommended in the posterior region due to their relatively high fracture rate compared to titanium.120 Still, most published clinical trials provide data on 5 years or less of in vivo performance. It is uncertain what the success rate would be over a more extended observation period given its low fracture toughness and tendency for degradation.

Milling of metal structures, rather than using the traditional lost wax technique, allows for more predictable and less technique-sensitive production of metal frameworks for dental prostheses. There are two types of metals commonly milled: titanium and cobalt-chromium. Milled titanium implant abutments and frameworks have become more popular as they are deemed to combine the best features of traditional techniques. Titanium has high accuracy of the stock abutment and yet is modifiable through the CAD phase so that it does not require modification after milling to achieve the desired dimensions.115 Titanium full-arch implant frameworks and custom abutments are reported to have excellent success.120,121 In the case of cobalt-chromium structures, base metals are generally challenging to cast due to shrinkage during the solidification phase and therefore have a higher risk of distortion and inaccuracy.122 Subtractive manufacturing does not usually involve elevated temperatures, reducing the risk of distortion. However, the hardness of the cobalt-chromium does significantly increase the wear of burs when milled in its fully sintered state using hard machining processes.123 The use of soft milling/machining of presintered cobalt-chromium blocks increases the life of the milling burs, although the sintering phase of the product post-milling induces shrinkage of the structure that has to be accounted for prior to fabrication.124 This manufacturing method also leads to differences in structure and behavior of produced cobalt-chromium alloy compared to conventionally cast cobalt-chromium, the milled alloy. Milled cobalt-chromium, whether hard or soft machined, has a more homogenous microstructure that is mostly free from critical defects when compared to conventionally cast cobalt-chromium. It has a relatively small grain size with regular grain boundaries, and this is associated with improved mechanical properties of the overall alloy.125 In vitro studies have also shown that milled cobalt-chromium has much improved bonding to porcelain and mechanical properties than its conventionally cast counterpart.126,127 However, it has also been shown that soft- machined cobalt-chromium results in slightly inferior bonding to porcelain and mechanical properties when compared to hard-machined cobalt-chromium. This is due to the inherent porosities formed during the sintering step and relatively lower ductility, although both have mechanical properties that exceed required standards for dental prostheses.127

Subtractive manufacturing methods have also been found to be highly accurate. A systematic review of marginal accuracy of crowns produced using subtractive technology found that all measurements of marginal gaps were below 120 μm and most were around 80 μm, which is well within acceptable marginal fit discrepancy tolerance. However, internal gaps varied from 34 μm to 220 μm in the occlusal fitting surface depending on the area of measurement.128 For manufacturing implant frameworks, the fit was found to be excellent, with less than 4 μm vertical discrepancy and less than 15 μm in a passive fit. Similarly, for custom-made abutments, the accuracy was found to be 2.5 μm to 3.2 μm, with rotational movement less than 3 degrees.129

Subtractive manufacturing techniques do have an issue with high wastage. Some studies have reported that up to 90% of the starting material is lost when machining of dental prostheses from a raw block.130 Although excess material is usually unrecyclable after use, subtractive manufacturing technique demonstrates very low efficiency and is undesirable in cases where the raw material cost is an issue. For example, the amount of wastage involved in milling high-priced noble metals, such as palladium or gold, would be very prohibitive, and hence subtractive manufacturing would not be an economically viable option for use on such materials.131 Finally, subtractive manufacturing is not feasible when the desired part is larger than the preformed blocks and disks or when the 3D geometry is too complex even with operator reindexing on the most versatile milling systems.

Printed materials

The alternative to subtractive manufacturing in the CAM step of the dental workflow is additive manufacturing technique (3D printing). In 1986, Charles Hull introduced the principle of additive manufacturing. ASTM International, formerly known as the American Society for Testing and Materials, defines additive manufacturing as “the process of joining materials to make objects from 3D model data, usually layer upon layer, as opposed to subtractive manufacturing methodologies.”131 This manufacturing method has been extensively used as an efficient method for rapid prototyping when highly customized models are required,132 thus making it very suitable for the highly individualized prosthesis that is required in clinical dentistry.

In subtractive manufacturing, CAD data is converted into CAM files to program the tooling paths for cutting tips. In contrast, most 3D printers convert the CAD data into thin vertical slices and program the deposition of materials or energy to produce parts. The slicing allows for input of infinitely complex geometries including internal cavities; closed-end, one-way curved channels133 such as those required in case planning; and metal plate fabrication for maxillofacial surgeries. Another point of difference is the passive nature of this manufacturing technique when compared to milling, especially with hard-to-machine metals such as cobalt-chromium, that results in wear of the milling heads of subtractive units, noise, and heat production during milling as well as surface damage of the structure.134

Additive machines, in contrast to milling machines, tend to have higher material efficiency, with approximately 40% less wastage. In addition, around 95% to 98% of the waste may be recycled in future production cycles.135,136 This can significantly reduce the overall cost of the raw material and is particularly crucial in situations where the overall weight and size of the raw material is an issue.

Different methods fall under the term additive manufacturing; however, the commonly used method in the manufacture of implant prostheses involves selective laser sintering (SLS) and selective laser melting (SLM). The basic premise is that a layer of powder (alloy or ceramic) is spread out and the particles are joined together by a focused laser with a preprogramed path. The support table then lowers, another layer of powder is spread, and the process repeats until the whole structure is created. While early application of technology produced very rough and porous structures that required extensive post-processing, more modern laser sintering machines use a chamber with no oxygen and have honed the former technique much further, with a tightly controlled sintering environment and power levels so the produced structure requires little post-processing.137 This production method is highly useful in dentistry due to the density of metals produced, the minimal post-processing required, and the complex geometries that may be involved in the fabrication of metal-based prostheses.138 Examples include PXM (Phenix Systems) and EOSINT M 280/M 270 (EOS).

There are a variety of material options used for additive manufacturing techniques. Titanium fabricated using SLM techniques has been shown to have high structural homogeneity and superior mechanical properties compared to SM or conventionally cast titanium.139 Additive manufacturing using SLM techniques is a promising method of cobalt- chromium production as well (Fig 4-8). The high hardness and mechanical properties of cobalt-chromium make the passive nature of 3D printing a much more efficient process with reduced stress on components and less energy consumption and wastage.134 In the past, laser sintering–based machines produced rather weak and porous structures that required extensive post-processing.140 A variation of the technique under vacuum has shown much improved results, producing dense end products.141 The resultant cobalt-chromium has been shown to have a uniform microstructure that is stable even after being subjected to high temperatures during porcelain firing cycles. It has also been shown to have significantly higher mechanical properties and stronger bonding to overlaying porcelain compared to conventionally cast or soft-machined manufactured cobalt-chromium but on par with hard-machined cobalt-chromium.142

Fig 4-8 Selective laser melting was used to fabricate this removable partial denture framework.

The accuracy of additively manufactured structures varies depending on the process and material. 3D-printed dental casts varied in accuracy depending on the type of process involved, with digital light processing shown to have higher accuracy than stereolithography methods, but both are considered superior to fused deposition modeling. While 3D-printed dental casts generally have excellent accuracy for single-crown fabrications,143 accuracy of casts used for fabrication of complete-arch implant prostheses is still considered inferior to an accurately recorded dental stone cast, although both are thought to be clinically satisfactory.144 Studies comparing the accuracy of PFM crowns produced using new SLM technology have been found to have acceptable marginal misfit for use for a dental prosthesis. For the manufacture of implant abutments, the process has been reported to leave an 11-μm vertical gap between the prosthesis and fixture and hence is still inferior to a milled implant abutment.115 The use of manufacturer-specific titanium bases (Ti-bases) as an interface between the prosthesis and fixture can therefore be used to ensure a more accurate fit.

While 3D printing has many applications and material options, it still has limitations. One example is the use of additive manufacturing for the fabrication of ceramic structures (zirconia and alumina) that results in porous structures, requiring extensive post-processing and in turn shrinkage. For this reason, 3D printing does not at this time have the uniformity that subtractive manufacturing produces and does not bypass the problem of shrinkage when machining presintered blocks of material.145 Previous limitations of 3D printing include slow print speed and “staircase”-like finish due to thick layering processes that require extensive polishing before use.129 Advanced printers today have mostly eliminated the staircase effect, and layerless printing is reducing print time from hours to minutes. Volumetric 3D printers that build entire 3D parts at once will open up an entirely new world for additive manufacturing, and these exciting developments will motivate more innovative materials that will benefit dentistry.

Occlusal materials

Selection of materials to withstand occlusal forces is often based on material strength measured in vitro, which may not always predict the true performance characteristics of a material when used in vivo. Dental restorations may be subjected to significant loads in the mouth reported in Newtons, particularly due to parafunctional activity,146–148 and small contact points between two opposing teeth may dramatically increase the stress that a restoration undergoes when a maximum bite force is encountered. This may be one of the reasons for the disparity between the restorative material’s strength values and clinical survival rates. One needs to keep in mind that strength is measured as a stress value, usually in MPa, based on the force in Newtons divided by the surface area to which the force was subjected. Therefore, the reporting of bite forces is not meaningful until the surface area to which the force was applied is known, usually including the occlusal contact points. A study by Jansen van Vuuren et al149 reported the bite forces in relation to contact points between opposing teeth in order to describe the stress that the teeth and future restorations on those teeth may experience. The study considered force and surface area and found that the resultant stress teeth may be subjected to potentially exceeded the flexural strength of even the strongest dental ceramic available on the market for 21.8% of the teeth tested. For this reason, clinicians need to carefully consider the material properties and strength values of the materials they select in treatment planning for posterior restorations.

Type III and IV noble alloys

Gold is one of the oldest materials used in prosthodontics. Its popularity arose due to its excellent corrosion and tarnish resistance, its good ductility, and its ability to be work hardened, making it an ideal candidate for use in various dental applications. However, due to the rising cost of gold and its limited esthetics as well as the strength compared to alternative materials, its usage today has been reduced and has been replaced by base metal alloys and high-strength ceramics.

Traditionally, dental gold alloys are categorized from type I to type IV according to their strength (low to high). Type I and II gold alloys are limited to inlays and onlays, type III gold is used for high-stress applications such as full crowns, and type IV gold is used for very high-stress applications such as multi-unit prostheses, bars, and partial denture frameworks. Because pure gold by itself is relatively weak for most applications, the high-strength variants (types III and IV) necessitate a certain amount of other elements (eg, Pd, Sn, Cu, Co, Pt, Ni, Co, Zn) to be incorporated in order to improve the mechanical properties of the alloy. As an occlusal material, single-unit gold alloy crowns have been shown to have a high survival rate of 96% at 10 years,150 which has been considered the “gold standard” for most other materials (Fig 4-9). Even for more esthetic restorations where a gold alloy is used as a substructure for a PFM crown, there are similarly high survival rates of 95.5% to 97.6% at 10 years.151–153

Fig 4-9 (a and b) Noble alloys remain the most predictable occlusal material for posterior implant-supported prostheses. The custom abutments were milled of titanium.

Lithium disilicate

First introduced in 1988 as a pressable ceramic, lithium disilicate is a crystalline filler used in particle-filled ceramics under the product names IPS e.max CAD and IPS e.max Press sold by Ivoclar Vivadent. In its final form, the lithium disilicate crystals take up approximately 70% of the material volume with the remainder being amorphous glass. For IPS e.max CAD, which is designed for milling, the material comes in a partially sintered lithium metasilicate form that has weak mechanical properties for easier milling and reduced wear on the burs before undergoing a crystallization fire at approximately 850°C to transform the ceramic into lithium disilicate. Due to its high flexural strength (~400 MPa) and fracture toughness (~2.5 MPa • m½)154,155 and acceptable esthetics compared to traditional leucite reinforced ceramics (< 150 MPa), lithium disilicate has gained considerable popularity for use in a wide range of applications, including implant-supported fixed dental prostheses (Fig 4-10). More recently, other companies have introduced their own particle- filled ceramic using zirconia-toughened lithium silicate, which has been shown to perform similarly. Clinically, lithium disilicate has been shown to have a high survival rate on par with the “gold standard” for single-unit restorations. For three-unit fixed dental prostheses, the 10-year survival rate is still acceptable at 71.4% to 87.9%.78,156

Fig 4-10 (a and b) Lithium disilicate has become increasingly popular for restoring implant-supported fixed dental prostheses, especially when the restoration extends into the esthetic zone. (Courtesy of Dr A. Pozzi.)

3Y-TZP and monolithic zirconia

First introduced around 2002 as a strong substructure ceramic that was an alternative to the less esthetic base or noble alloys, Y-TZP ceramics could be layered with a compatible porcelain or even cemented as a substrate material with the lithium disilicate ceramic. However, controversy surrounding the LTD issue due to the catastrophic fracture of femoral heads in orthopedics157 was a cautionary tale for dentistry. For the next decade after its release, much research was done to identify whether the role of LTD had a significant detriment to the clinical applications in dentistry. A plethora of research was able to identify the occurrence of LTD through simulated in vitro oral conditions and subsequent reduction in mechanical properties.51,158–161 However, no study was able to conclusively link the effect of LTD to the clinical survivability of Y-TZP ceramics. Nevertheless, this also led to research in alternative dopants such as cerium and germanium,162–164 where some had better resistance to LTD, though little was explored beyond that due to a more troubling issue: the chipping and fracture of veneers layered on Y-TZP. More importantly, the femoral zirconia head catastrophe highlights the importance for precision control over processing conditions. Dental laboratories should calibrate their ceramic furnaces routinely to ensure proper temperature control and follow best practices to control the cooling rate to minimize un- desirable phase transformations and microstructure changes. If the actual oven temperature is below the sensor reading, the suboptimal temperature may not be adequate to fully “heal” heterogeneous defects during puck manufacturing (powder compaction, presintering) and CAM/CAM (milling, post- processing, sintering). Conversely, if the actual oven temperature is higher than programming temperature, variable phase transformation can be expected. More importantly, the higher temperature will also cool faster when the prosthesis is open to room air due to greater thermal gradient. Faster cooling rates are often associated with higher internal residual stresses, especially in structures with irregular surface area–to-volume ratio due to 3D shape complexity. Ovens should be cleaned regularly, and pollution-absorbent powders should be used to reduce contamination. Thermocouples should be replaced if the oven uses silicon-containing heating elements (eg, MoSi₂) that leave a quartz-like layer on the thermocouple.

Besides bulk defects, another problem was the increased incidence of veneer chipping among Y-TZP restorations. Fortunately, it was soon discovered that the origin of the cause was the poor thermal conductivity and diffusivity of Y-TZP, resulting in a large buildup of residual stresses close to the interface of the bilayer restoration.165 A solution was found by slowing the cooling cycle performed on the manufacturing of these bilayered restorations, which dramatically reduced the residual stresses built up within the restoration. However, the limitation of the mechanical properties of the veneer material was still a deciding factor on the overall strength of these restorations in many studies that continued to report the issue of veneer fractures in bilayered Y-TZP.166,167 This led to the release of monolithic Y-TZP restorations.

The use of monolithic zirconia restorations (Fig 4-11) over its bilayer application eliminated several key problems: debonding and strength limitation of the veneer layer. Monolithic zirconia curently comes in several different forms tailored toward a balance between strength and esthetics. For more esthetic highly translucent forms of zirconia, there is a larger percentage of cubic phase present that requires a higher percentage of yttria to stabilize the crystalline structure. These options allow clinicians to select the translucency of choice suitable for the desired situation and placement. One of the main concerns of using monolithic zirconia is the increased wear on the antagonist dentition due to the high hardness of Y-TZP–based restorations. However, a number of current studies have found that monolithic zirconia actually results in less wear168,169 on the antagonist enamel compared to alternate materials due to their highly smooth surface that does not roughen easily with time. Little is currently available on the long-term survival rate of this still relatively new material, but the reported short-term success is promising.170

Fig 4-11 (a to c) Monolithic zirconia has recently been promoted as a suitable occlusal material when fabricating implant-supported xed dental prostheses.

Hybrid ceramics

One alternative approach to dental ceramics is known as the hybrid ceramic. These ceramics consist of either a porous ceramic structure or highly dense ceramic particles bounded by a highly crosslinked organic matrix/network. The essence of the technology is to produce interpenetrating phase composites, and the controlled manufacturing process allows a much more refined product. In terms of mechanical properties, hybrid ceramics are generally stronger than traditional feldspar- or glass-based ceramics, with flexural strength values between 150 and 200 MPa and fracture toughness of approximately 1.7 MPa • m½.171 One of the key advantages of this new hybrid ceramic material is the milling efficiency in both speed and wear,172,173 and the marginal accuracy is much better compared to other machinable ceramic blocks (eg, e.max and feldspar-based ceramics). The wear characteristics of some hybrid ceramics such as the Vita Enamic (Vita Zahnfabrik) has also been found to be similar to that of enamel.174 Being a relatively new material, little information is available on the clinical survival rates, but the mechanical properties would indicate that they should perform acceptably within their indicated use of single-unit restorations.

Denture teeth

The selection of denture teeth for implant overdentures and hybrid implant dentures is not straightforward. The forces applied to the teeth and denture base resins of implant overdentures are of a much higher magnitude than that of conventional mucosa-borne dentures due to reduced mobility of implant-borne dentures and the lack of proprioceptive feedback around implants. Hence, the combination of higher forces and constrained movement results in frequent maintenance issues such as dislodgment and fracture of denture teeth.175 While porcelain teeth have excellent esthetics and wear resistance, there is no chemical bond between the teeth and the denture base resin. They are mostly dependent on the mechanical retention tags in the anterior teeth and the hollowed-out shape on the underside in the posterior teeth. In closed bite situations, this is often reduced by grinding to fit them into the limited space. Therefore, the selection of porcelain teeth presents a high risk of dislodgment.

Composite denture teeth have very good esthetics and wear properties but are limited in the amount of chemical bonding that can be achieved if the PMMA layer on the underside of the ridge-lap component of the teeth is ground off while setting up the teeth in closed bite or limited space situations. Once the resin layer is removed, the high-percentage ceramic filler component is exposed, which does not bond chemically to the denture base resin; only the resin matrix will bond. Therefore, the selection of composite teeth presents a high risk of dislodgment if the resin layer has been ground off. High-quality PMMA teeth have good esthetics but are not as wear resistant as porcelain and composite teeth (Fig 4-12). Composite teeth are more resistant to wear but are at risk for chipping and fracture, especially when the screw access channel exits through the occlusal surfaces (Fig 4-13). However, a good chemical bond to the denture base resins is achieved, and these types of teeth are less likely to dislodge from the dentures in comparison to porcelain and composite teeth. Finally, denture teeth around the prosthetic screw channels of hybrid prostheses are prone to dislodgement, wear, and fracture, regardless of teeth materials (see chapter 20, section entitled, “Fracture and chipping of denture teeth”).

Fig 4-12 Acrylic resin denture teeth are subject to rapid wear when used in implant-retained and -supported prostheses.

Fig 4-13 When the occlusal access channel exits through the occlusal surfaces of composite denture teeth, they are at risk of chipping and fracture.

Summary

Despite their documented clinical success, none of the materials presented in this chapter are fully resistant to the chemical, biochemical, mechanical, and biologic factors in the local oral implant microenvironment. All materials presented undergo varying levels of intraoral breakdown by degradation, erosion, corrosion, and wear, and some of the materials produce breakdown debris that can cause cytotoxicity and immune reactivity. As the search continues for the ideal materials for each major component within the implant-prosthesis system (implant fixtures, prosthesis, substructure, bars, and teeth), some of the fundamental science presented in this chapter may provide the rationale for understanding the role of the future dopants and additives. It is worth noting that although a disproportionately large fraction of this chapter is related to zirconia, it is not intended to suggest endorsement of it. In fact, the authors have observed clinical fractures, chipping, and peri-implantitis and therefore recommend caution with this material—especially its use as dental implant fixtures. It is exciting to reflect upon the fact that today’s materials for implant dentistry are much better than our repertoire from only 10 years ago, and much more will come as the deployment of machine learning in modern materials engineering and polymer synthesis will accelerate the pace of advanced materials development.

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