**2. Yttria-partially stabilized tetragonal zirconia polycrystals (Y-TZP)**

#### **2.1. Material features**

experimental conditions tested. The proliferation rate was consistent with that expect‐ ed for biocompatible materials. Hydrothermally treated ATZ samples and Si3N4 rough surfaces were capable to enhance the osteogenesis in vitro. The biological responses induced in MC3T3 cells were correlated with the surface features. Immediately after seeded, osteoblasts are known to interact with their substrate via integrins that bind to the proteins adsorbed on the biomaterial surface. The interface effect was discussed in light of the literature. The most recent publications suggest that research aims at investigating the effects of surface modifications dictating the chemical characteristics and the nano-/micro-topography that are paramount modulators of the biological

**Keywords:** surface roughness, dental implants, ceramic materials, surface modifica‐

Modern oral implantology has been based on titanium since the research line originated by Brånemark's first discovery and subsequent experiments [1]. Titanium implants have the longest traceable record of predictable clinical performance with a cumulative success rate of 98.8% for 15 years [2]. High biocompatibility, favorable tissue response and adequate strength and corrosion resistance rendered titanium implants widely diffused in the market. The number of dental implant brands grew from 45 systems in 1988 [3] to 600 systems produced by 146 manufacturers in 2008 [4]. Currently, worldwide, there are more than 350 dental implant manufacturers producing an estimate of 1600 different systems, 98% of which are titanium implants. Titanium, however, is no longer considered a completely bioinert material, instead it might be an allergen as reported by several studies [5–8]. Elevated titanium concentrations have been found in the vicinity of oral implants [8], in regional lymph nodes [9], serum and urine [10], which is potentially hazardous to human body. Besides these issues, some dental patients are metal-phobic and demand to be treated solely with metal-free dental implants [11]. Only recently, truly viable alternative materials were proposed to titanium, although the first ceramics for oral applications dated back to the 1970s. Historically, indeed, high-density, highpurity aluminum oxide (alumina) was chosen for dental implant manufacturing, as it com‐ bined excellent corrosion resistance, good bio-compatibility, high wear resistance, and high strength. Despite these promising features, the material was brittle and prone to fracture under unfavorable load. Thus, the positive preclinical and clinical outcomes of the first studies could not prevent alumina implant systems to be withdrawn from the market [12]. Research and manufacturing technology have greatly improved the offer of bio-ceramics, thanks to the introduction of yttria-partially stabilized tetragonal zirconia polycrystals (Y-TZP), whilst a possible future use of alumina zirconia composites and silicon nitride–titanium nitride composites may further expand the offer of reliable devices on the market. Three distinct sections of the present chapter are dedicated to each of these materials. Specifically, the literature regarding zirconia was thoroughly revised in Section 2, whilst some novel data of our group are exposed and discussed in light of and along with previous work as for alumina

response.

64 Dental Implantology and Biomaterial

**1. Introduction**

tions, interface

Zirconia (ZrO2) is a crystalline dioxide of zirconium: as thoroughly reviewed elsewhere [12], unalloyed zirconia can assume three crystallographic forms depending on the temperature, at ambient pressure. At room temperature and upon heating up to 1170°C, the symmetry is monoclinic (P21/c). The structure becomes tetragonal (P42/nmc) between 1170 and 2370°C and cubic (Fm3m) above 2370°C and up to the melting point [13]. Upon cooling at ~950°C, during the transformation from the tetragonal (t) phase to the monoclinic (m) phase, a substantial increase in volume (~4.5%) occurs, which is sufficient to lead to catastrophic failure. By alloying pure zirconia with stabilizers such as calcium oxide (CaO), magnesium oxide (MgO), yttrium oxide (Y2O3), or cerium oxide (CeO2), the tetragonal structure is maintained, even at room temperature, and the stress-induced t → m transformation is controlled, efficiently arresting crack propagation [14, 15]. Indeed, when a crack develops, tetragonal grains convert imme‐ diately to monoclinic form. The propagation of the crack develops sufficient stress within the tetragonal structure to transform also the grains around the crack to stable monoclinic form. Thus, the expansion volume of zirconium dioxide crystals produces compressive stress around the crack and prevents further propagation of crack [16–18]. This mechanism is known as transformation toughening and is influenced by temperature, vapor, particle size, micro- and macrostructure, and concentration of stabilizing oxides [19].

Yttria-stabilized zirconia (Y-TZP) [20] is endowed with excellent mechanical, and tribological properties together with biocompatibility and rightly represent a good choice for preparing dental implants. As yttria decreases the driving force of the t-m transformation [21, 22], biomedical grade zirconia are usually stabilized with 3 mol% yttria (Y2O3) (hence 3Y-TZP) [16]. The salient mechanical properties of Y-TZP are reported in **Table 1**, but it is noteworthy that the Weibull modulus is strongly dependent on the type of surface finish and the processing conditions [23].

Notwithstanding the excellent mechanical properties of Y-TZP [19, 24], recent reports of in vivo failures [25–27] have questioned the long-term stability of the material. The low-temper‐ ature degradation (LTD) of zirconia [22, 28–31], also known as aging process, plays here a fundamental role. Involving the t→m transformation, LTD can be favored, even at room temperature, by the penetration of water radicals into zirconia lattice, thus leading to the formation of tensile stresses in zirconia surfaces. The activation barrier for the transformation is lowered, and the phase transition is promoted. The main consequences of this aging process include surface degradation with grain pullout, microcracking, and strength degradation. As reported by Cattani-Lorente et al. [32] also, Young's modulus and hardness of Y-TZP bars were reduced by 30%, when they were subjected to hydrothermal cycling. The increase of mono‐ clinic–tetragonal phase ratio was associated with microcracking and resulted responsible for the decline in mechanical parameters [32].

To control the aging phenomenon, several factors can be taken into account: from the obvious use of stabilizers to the modulation of residual stress [33]. Likewise, adjusting crystal size and removing impurities during manufacturing was proposed with the same anti-aging scope [34]. Interestingly, surface finishing could affect the aging kinetics of 3Y-TZP, according to Deville et al. [35]. More precisely, rough polishing produced a compressive surface stress layer beneficial for the aging resistance, whilst smooth polishing lead preferential transformation nucleation around scratches, due to elastic/plastic damage tensile residual stresses.

In an extensive review of his, Jerome Chevalier concluded that "although in the 1990s, 3Y-TZP ceramics were considered very promising materials for biomedical applications, long-term follow-up is needed to address the critical problem of aging in vivo. Moreover, most zirconia implants were processed at a time when aging was not yet fully understood. Methods to assess a priori the aging sensitivity of a given zirconia ceramic have been developed and should lead to safer implants. In the meantime, new zirconia or zirconia-based materials that overcome the major drawback of the standard 3Y-TZP are now available" [22].

#### **2.2. Manufacturing methods**

Hot isostatic press (HIP) is the most common method used for preparing zirconia dental implant. By subjecting encapsulated powder, or sintered yet porous parts, to inert gas at isostatic pressure at a high temperature, HIPing is deemed an excellent method to obtain highdensity homogenous products [36]. HIPing enables the application of an equally distributed pressure in all directions resulting in greater material uniformity and higher strength [2]. HIPing of Y-TZP enhances the strength, eliminates fracture sources such as pores, and reduces the aging phenomenon [37]. The preparation entails many steps as summarized in **Figure 1**. Briefly, Y-TZP blocks are presintered at temperatures below 1500°C to reach a density of at least 95% of the theoretical density. Hot isostatic pressing (HIP) is applied to the blocks at temperatures between 1400 and 1500°C under high pressure. A HIP cycle after sintering is recommended to achieve a full density close to the theoretical values (d = 6.1 g/cm3 = 100% dense). Since HIPing changes the color of Y-TZP into dark-grey, a heat treatment in air is usually performed to restore the material whiteness by oxidation, prior to be machined using a specially designed milling system. Because of the high hardness of fully sintered Y-TZP, the milling system is to be particularly robust [38–40].

**Figure 1.** Manufacturing process for Y-ZPT.

The relatively recent and yet pervasive introduction of computer-aided design/computeraided manufacturing (CAD/CAM) technology has provided dentistry with an alternative to HIPing [42]. Usually, dental CAD/CAM systems recur to partly sintered yttria-stabilized tetragonal zirconia polycrystal (Y-TZP) blanks. The use of this partly sintered state of the Y-TZP ceramic renders the milling process faster and reduces the tools wear, compared to systems employing densely sintered blanks (HIP process). Of course, the final sintering shrinkage must be taken into account during the CAD phase by enlarging the shapes before milling, whilst this compensation is not necessary with the HIPed Y-TZP blanks that are directly ground to the desired dimensions [5]. The salient mechanical properties of Y-TZP subjected to the two manufacturing work-flow described above are compared in the following [43] table (**Table 1**), along with pressed and sintered polycrystalline α-alumina.


**Table 1.** Values refer to Duraccio et al. [12].

#### **2.3. Biological properties**

clinic–tetragonal phase ratio was associated with microcracking and resulted responsible for

To control the aging phenomenon, several factors can be taken into account: from the obvious use of stabilizers to the modulation of residual stress [33]. Likewise, adjusting crystal size and removing impurities during manufacturing was proposed with the same anti-aging scope [34]. Interestingly, surface finishing could affect the aging kinetics of 3Y-TZP, according to Deville et al. [35]. More precisely, rough polishing produced a compressive surface stress layer beneficial for the aging resistance, whilst smooth polishing lead preferential transformation

In an extensive review of his, Jerome Chevalier concluded that "although in the 1990s, 3Y-TZP ceramics were considered very promising materials for biomedical applications, long-term follow-up is needed to address the critical problem of aging in vivo. Moreover, most zirconia implants were processed at a time when aging was not yet fully understood. Methods to assess a priori the aging sensitivity of a given zirconia ceramic have been developed and should lead to safer implants. In the meantime, new zirconia or zirconia-based materials that overcome the

Hot isostatic press (HIP) is the most common method used for preparing zirconia dental implant. By subjecting encapsulated powder, or sintered yet porous parts, to inert gas at isostatic pressure at a high temperature, HIPing is deemed an excellent method to obtain highdensity homogenous products [36]. HIPing enables the application of an equally distributed pressure in all directions resulting in greater material uniformity and higher strength [2]. HIPing of Y-TZP enhances the strength, eliminates fracture sources such as pores, and reduces the aging phenomenon [37]. The preparation entails many steps as summarized in **Figure 1**. Briefly, Y-TZP blocks are presintered at temperatures below 1500°C to reach a density of at least 95% of the theoretical density. Hot isostatic pressing (HIP) is applied to the blocks at temperatures between 1400 and 1500°C under high pressure. A HIP cycle after sintering is

recommended to achieve a full density close to the theoretical values (d = 6.1 g/cm3

dense). Since HIPing changes the color of Y-TZP into dark-grey, a heat treatment in air is usually performed to restore the material whiteness by oxidation, prior to be machined using a specially designed milling system. Because of the high hardness of fully sintered Y-TZP, the

The relatively recent and yet pervasive introduction of computer-aided design/computeraided manufacturing (CAD/CAM) technology has provided dentistry with an alternative to HIPing [42]. Usually, dental CAD/CAM systems recur to partly sintered yttria-stabilized

= 100%

nucleation around scratches, due to elastic/plastic damage tensile residual stresses.

major drawback of the standard 3Y-TZP are now available" [22].

milling system is to be particularly robust [38–40].

**Figure 1.** Manufacturing process for Y-ZPT.

the decline in mechanical parameters [32].

66 Dental Implantology and Biomaterial

**2.2. Manufacturing methods**

In vitro experiments on different cell lines, in vivo studies on animals and clinical studies on humans supported the safety and the high level of biocompatibility of zirconia. In a prelimi‐ nary in vitro investigation [44], one-piece zirconia implants were proven to possibly fulfill the biomechanical requirements for anterior teeth restoration. In addition, the mean fracture strength of zirconia implants was investigated after chewing simulation and it was found to be within the limits of clinical acceptance. However, the preparation of a one-piece zirconia implant prior to prosthetic finalization may significantly compromise fracture strength. Therefore, long-term clinical data were deemed necessary before one-piece zirconia implants could be recommended for clinical practice [45]. For the same reason, two-piece zirconia implants were considered clinically inadequate due to the increased risk of fracture at the implant head level [46].

Evidence from in vitro studies on osteoblasts supported the possible favorable response of zirconia ceramics in vivo [47, 48]. When implanted in bone or soft tissues, these materials could elicit no inflammatory reactions, nor fibrous encapsulation, according to Hisbergues et al. [41]. Interestingly, Scarano et al. [49] reported the osseointegration of unloaded zirconia implants inserted in rabbit bones without any signs of inflammation or mobility. The possible role of surface roughness was investigated by comparing the removal torque of machined zirconia implants to roughened ones [50]. Notably, the roughened implants performed better than the smooth ones and behaved similarly to the oxidized titanium implants used as control. Loaded zirconia implants were studied and compared to titanium implants by Kohal and co-workers [51], who could find no significant difference in the osseointegration level between the two groups. Akagwa et al. [52] reported a similar bone to implant interlock in loaded and unloaded zirconia implants, but a crestal bone loss higher around the former group. In favor of the clinical use of Zirconia, it must be cited its maintenance of bending strength of over 700 MPa after immersion in 95°C saline solution for over 3 years [53]. Furthermore, zirconia blanks did not show any significant mechanical detriment even after being embedded in the medullary cavity of the tibia of rabbits for 30 months.

Scarce are the clinical studies dedicated to the long-term performance of zirconia implants. The short follow-up period and the often small sample size hinder their quality of evidence, so that Andreiotelli and coauthors [29] could only include three retrospective cohort studies on one-piece zirconia dental implants in their systematic review, reaching in total 231 patients and 416 implants. The studies by Mellinghoff et al. [54] and Oliva et al. [55] investigated, respectively, 189 and 100 zirconia implants and estimated 1-year survival rates of 93 and 98%. Almost all of the failures occurred during the healing phase, as only one implant failed after prosthetic reconstruction due to fracture. Lambrich and Iglhaut [56] observed 127 zirconia and 234 titanium implants for a mean period of 21.4 months. Notably, in this study, the survival rate of zirconia implants was similar to that of titanium in the mandible (Y-TZP = 98.4% vs. Ti = 97.2%), whilst differed considerably in the maxilla (Ti = 98.4% vs. Y-TZP = 84.4%). Again, all failures occurred during the healing phase owing to increased implant mobility. These findings are consistent with the paper by Depprich et al. [57], where the survival rate of zirconia implants obtained from 17 clinical studies was between 74 and 98% after 12–56 months. Payer et al. [58] followed up for 2 years 19 immediately loaded zirconia implants, reporting a 95% survival rate, as determined clinically and radiographically. These results are in accordance with Oliva et al. [59] who determined the same survival rate at 5 years in 371 patients who received 831 one-piece zirconia implants. Kohal et al. [60] found that immediately restored one-piece zirconia implants have 1-year cumulative survival rate comparable to titanium counterparts. In conclusion, the clinical data currently available for Y-TZP implants may not be sufficient to recommend their routine clinical use. Zirconia, however, may have the potential to be a successful implant material, although this is as yet not fully supported by present investigations and further good-quality research is needed.
