Zirconia based ceramics, some clinical and biological aspects: Review Academic research paper on "Medical engineering"

Accepted Manuscript

Zirconia based ceramics, some clinical and biological aspects: Review Ossama Saleh Abd El-Ghany, Ashraf Husein Sherief

PII: S2314-7180(16)30039-8

DOI: 10.1016/j.fdj.2016.10.002

Reference: FDJ 19

To appear in: Future Dental Journal

Received Date: 22 August 2016 Accepted Date: 3 October 2016

Please cite this article as: Abd El-Ghany OS, Sherief AH, Zirconia based ceramics, some clinical and biological aspects: Review, Future Dental Journal (2016), doi: 10.1016/j.fdj.2016.10.002.

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DENTAL^

JOURNAL

Zirconia Based Ceramics, Some Clinical and Biological Aspects: Review

Ossama Saleh Abd El-Ghanya , Ashraf Husein Sherief

a Professor, Department of Fixed Prosthodontics, Faculty of Oral and Dental Medicine, Future University in Egypt FUE

b Professor, Department of Fixed Prosthodontics, Faculty of Oral and Dental Medicine, Future University in Egypt FUE

Corresponding author

E-mail address: osama.abdelghany@fue.edu.eg.com (Osama Saleh Abd El-Ghany).

Key words

zirconia ceramics, machining, bonding, esthetics, biological aspect, radioactivity. Abstract

Improved material strength, enhanced esthethic and high biocompatibility give Zirconia ceramic a great possibility to be used for a wide range of promising clinical applications. This review presents the different types of zirconia materials available for dental application, the effect of machining procedures on these materials, the esthetic of zirconia ceramics and bonding of the veneering ceramics in addition to the biologic properties of these new materials.

Introduction

The Stone Age and Bronze Age were named for the materials that dominated these major historical periods. Consequently, the modern era, which is encountering an ever- increasing assortment of ceramic materials for industrial or biomedical use, could be characterized as the Ceramic Age. Ceramic materials that specially developed for medical and dental use are termed bioceramics. Over the last decade, zirconia technology has encouraged a rapid development of metal free dentistry that may provide high biocompatibility, enhanced esthetics and improved strength [1].

The name "Zirconium" comes from Arabic word "Zargon" which means "golden in colour" which in turn comes from the two Persian words Zar (gold) and Gun (colour). Zirconium dioxide (ZrO2) was accidentally dicovered by a German chemist his name is Martin Heinrich Klaproth in 1789 while he was working with certain procedures that involved the heating of some gems [2]. Subsequently, Zirconium dioxide was used as rare pigment for a long time. In late sixties the progress of research, the use of zirconium as biomaterials was refined. The first use of zirconium oxide (ZrO2) for medical purposes was made in 1969 in orthopedic as a new material for hip head replacement instead of titanium or alumina prostheses [3].

Zirconium (symbol Zr) is a transition metal element, atomic number 40,

atomic weight 91.22, density of 6.49g/cm , a melting point of 2,128K (1,855°C or 3,371°F) and a boiling of 4,682K (4,409°C or 7,968°F) and lustrous with exceptional corrosion resistance. Pure zirconium exists in crystalline form as a white and ductile metal and in an amorphous form as a blue black powder. Zirconium is ranked 18th in abundance among the element in earth's crust , however, this element does not occur in nature in a pure state but only in conjugation with silicate oxides (ZrO2 XSiO2) or as a free oxide (ZrO2) [4-7]. Zirconium Dioxide (ZrO2) is a white crystalline oxide of zirconium found in the minerals baddeleyite (ZrO2 ) and zircon (ZrO2).

Zirconium oxide crystals can be categorized into three crystallographic phases: the cubic phase (C) in the form of straight prism with square side, the tetragonal phase (T) in the form of a straight prism with rectangular sides, and the monoclinic phase (M) in the form of a deformed prism with parallelepiped sides. The cubic phase is stable above 2,370°C and with moderate mechanical properties, the tetragonal phase is stable between 1,170°C and 2,370°C with improved mechanical properties, and the monoclinic phase, which is stable at room temperature up to 1,170°C, with lower mechanical properties and may contribute to a reduction of the ceramic particles cohesion [8-10].

These lattice transformations are martensitic, characterized by being diffusionless (i.e. involving only coordinated shifts in lattice positions versus transport of atoms), occurring athermally implying the need for a temperature change over a range rather than at a specific temperature and, involving a shape deformation [4].This transformation range is bounded by the martensitic start (Ms) and martensitic finish temperatures. Volume changes on cooling associated with these transformations are very crucial and substantial so it results in a pure material that is unsuitable for any applications requiring an intact solid structure. This change is about 2.31% in case of C^T transformation and approximately 4.5% on cooling from T to M [5]. Phase transformation:

At ambient pressure, unalloyed zirconia can be found in three crystallographic forms depending on the temperature. At room temperature and upon heating up to 1170 °C, it is monoclinic. At temperature between 1170 and 2370 °C it is tetragonal and above 2370 °C and up to the melting point it is cubic. Upon cooling, the transformation from the tetragonal (t) phase to the monoclinic (m) phase is characterized by a substantial increase in volume (about 4.5%), sufficient to lead to catastrophic failure, sufficient to lead to catastrophic failure. This transformation begins at 950 °C and is reversible [11].

Garvie et al in their important paper" Ceramic Steel" explained the best method to use phase transformation of zirconia to improve the mechanical strength and toughness of this type of ceramics. They stated that tetragonal phase precipitates that are finely dispersed within the cubic matrix were able to be transformed into the monoclinic phase when the constraint exerted on them by the matrix was relieved by a crack propagated within the matrix. The stresses resulted from this phase transformation (into a larger size phase) will act in opposite direction to the stress fields that initiate and promotes crack propagation. Such mechanism will result in enhancement in fracture toughness as the energy associated with crack propagation is dissipated both in both phase transformation and in overcoming stresses due to the volume expansion.This process is known as transformation toughening [6].

It was discovered that such strong ceramic could result from alloying of pure zirconia with lower valance oxides, such as MgO, La2O3, CaO and Y2O3, and this decreased the amount of strained (m) phase at room temperature and favored more symmetric (c) and (t) lattice structures. These (c ) and (t) phases are analogous to those in pure zirconia but have dopant ions substituted on Zr4+ sites and have a fraction of oxygen sites vacant to retain charge neutrality[7].

Different types of zirconia ceramics materials available for dental applications:

Yttrium Tetragonal Zirconia Polycrystals (3Y-TZP)

Since the late eighties, biomedical grade zirconia usually contains 3mol% yttria (Y2O3) as a stabilizer (3Y-TZP) [8]. 3Y-TZP has been used to manufacture femoral heads in total hip replacement prostheses, but its use in orthopedic surgery has been reduced by more than 90%, mostly due to a series of failures that occurred in 2001[2]. Also 3Y-TZP is available in dentistry for the fabrication of dental crowns and fixed partial dentures. It has superior mechanical properties as its flexural strength reaches 900-1200 MPa and fracture strength about 9-10MPa(m)12 [12-13].

The restorations are processed either by soft machining of presintered blanks or by hard machining of fully sintered blocks. The mechanical properties of 3Y-TZP are strongly affected by its grain size [9]. Above a critical grain size, 3Y-TZP is less stable and more susceptible to spontaneous t^m transformation (grain sizes (<1^m) are associated with a lower transformation rate). In the same time, below a certain grain size about 0.2^m, the transformation is not possible, leading to reduced fracture toughness [10]. Consequently, the sintering conditions (through which the grain size is controlled) have a strong impact on both stability and mechanical properties of the final product. Higher sintering temperatures and

longer sintering times lead to larger grain sizes [11]. Currently available 3Y-TZP for soft machining of dental restorations utilizes final sintering temperatures between 1350 °C and 1550 ° C depending on the manufacturer.

Restorations produced by soft machining are sintered at a later stage (i.e. following the forming steps), this process prevents the stress-induced transformation from tetragonal to monoclinic and leads to a final surface virtually free of monoclinic phase unless grinding adjustments are needed or sandblasting is performed. Most manufacturers of 3Y-TZP blanks for dental applications do not recommend grinding or sandblasting to avoid both the t^m transformation and the formation of surface flaws that could be detrimental to the long-term performance, despite the apparent increase in strength due to the transformation-induced compressive stresses. In contrast, restorations produced by hard machining have been shown to contain a significant amount of monoclinic zirconia. This is usually associated with surface microcracks, higher susceptibility to low temperature degradation (LTD) and lower reliability [14].

Glass-Infiltrated Zirconia-Toughened Alumina (ZTA)

Another approach to advantageously utilize the stress induced transformation capability of zirconia is to combine it with an alumina matrix, leading to a Zirconia-Toughened Alumina (ZTA) [15]. One commercially available

dental product, In-Ceram Zirconia (VidentTM, Brea, CA), was developed by adding 33 vol.% of 12mol% ceria stabilized zirconia (12Ce-TZP) to In-Ceram Alumina [16]. In-Ceram Zirconia can be processed by either slip casting or soft machining. Initial sintering takes place at 1100 °C for 2h, followed by glass infiltration of this porous ceramic composite.

One of the main advantages of the slip-cast technique is the very limited amount of shrinkage. However, the amount of porosity is greater than that of sintered 3Y-TZP and comprises between 8 and 11% [17]. This explains the lower mechanical properties of In-Ceram Zirconia when compared to 3Y-TZP dental ceramics [12]. On the other hand Ce-TZP ceramics usually exhibit better thermal stability and more resistance to low temperature degradation than Y-TZP under similar aging conditions[18].

In-Ceram Zirconia for soft machining is thought to exhibit better mechanical properties compared to slip-cast ceramic due to more consistent processing. Conversely, Guazzato et al. reported a significantly higher flexural strength for In-Ceram Zirconia processed by slip-casting (630±58MPa) compared

to the soft machined material (476±50MPa) [19]. Also the fracture toughness varied

between 3.1 and 4.61 MPa(m) for both materials[20]. The two materials exhibited similar microstructure characterized by large alumina grains (6^m long, 2^m wide) together with clusters of small zirconia grains (less than 1^m in diameter).

The crack patterns were consistently transgranular for ZrO2 and intragranular for Al2O3. An advancing crack results in the tetragonal to monoclinic transformation. The associated increase in volume creates microcracks in the alumina matrix surrounding the transformed particle. The toughness is therefore enhanced by microcracking [21].

Magnesia Partially Stabilized Zirconia (Mg-PSZ)

Although a considerable amount of research has been focused on magnesia partially stabilized zirconia (Mg-PSZ) for possible biomedical applications, this material has not been successful. This mainly due to many causes as the presence of porosity, the large grain size (30-60^m) that can cause wear of the opposing structure, its low stability and low overall mechanical properties[22]. The microstructure consists of tetragonal precipitates within a cubic stabilized zirconia matrix. The amount of MgO in the composition of commercial materials usually ranges between 8 and 10mol%. In addition to a high sintering temperature (between 1680 and 1800 °C), the cooling cycle has to be strictly controlled, particularly in the aging stage with a preferred temperature of 1100 °C. Precipitation of the transformable t-phase occurs during this stage, which volume fraction is a critical factor in controlling the fracture toughness of the material [23].

It is very difficult to obtain Mg-PSZ precursors free of SiO2, so Magnesium Silicates can form and this decreases the Mg content in the grains and promote the

t^m transformation [24]. This can result in less stable material with lower mechanical properties. Denzir-M® (Dentronic AB) is an example of Mg-PSZ ceramic available for hard machining of dental restorations. Zirconia-Containing Lithium Silicate Ceramics (Zls)

Lithium silicate-based glass-ceramics were recently introduced as machinable materials to be used with (Computer aided design- computer aided manufacture (CAD-CAM) techniques, with claimed mechanical properties comparable with those of lithium disilicate glass ceramics. The technology for this new material relies on the addition of 10 wt% zirconium oxide to lithium silicate glass compositions. Zirconia crystals act as nucleating agent but remains in solution in the glassy matrix. A dual microstructure consisting of very fine lithium metasilicate (Li2SiO3) and lithium disilicate (Li2Si2O5) crystals is obtained, with a glassy matrix containing zirconium oxide in solution [25].

The microstructure is achieved through two stages. The glass-ceramic in the first pre-crystallized stage contains only lithium metasilicate crystals and is easy to machine. The final crystallization stage, leading to the dual lithium silicate microstructure, is obtained after a short heat treatment at 840°C for 8 min.

The main difference between ZLS and lithium disilicate glass ceramics glass ceramics in their final stage of crystallization resides in the nature of the crystalline phases: lithium metasilicate plus lithium disilicate for ZLS and lithium disilicate

only for lithium disilicate glass ceramic. The development of zirconia-containing lithium silicate glass ceramics illustrates the ongoing quest for ceramic materials that offer adequate translucency combined with superior mechanical properties. These stable ceramics may offer a better reliability than zirconia ceramics but may not represent the endpoint for this quest [25].

According to the manufacturers, these materials offer mechanical properties ranging from 370 to 420MPa. These values are comparable with lithium disilicate glass ceramics [26]. The mechanical properties are approximately three times higher than those determined for leucite-reinforced glass ceramics (IPS Empress, Ivoclar Vivadent, Schaan, Liechtenstein). In vitro testing revealed a Weibull modulus of 8.9 for this material group which indicates a uniform and reliable material quality. The improved strength and reliability are reached by the addition of 8-10wt% of zirconium oxide with a mean grit size of approximately 0.5 to 0.7 ^m [27].

Today, these materials are available as industrially prefabricated blanks for various CAD/CAM systems (e.g., Cerec, Sirona, Bensheim, Germany; Artica, KaVo, Leutkirchen, Germany; and Ceramill, Amann Girrbach, Pforzheim, Germany) in various shades and translucencies [28]. Rinke et al. stated that ZLS ceramics offer an excellent combination of high strength and outstanding optical properties. Thus, these materials are interesting for the fabrication of monolithic

restorations. However, although laboratory studies showed that ZLS ceramics have a positive combination of properties, the indication should be chosen with strict observation of the material-specific processing instructions. This attention to indications and processing instructions is especially important regarding the necessary minimum wall thickness and required adhesive luting. Long term clinical studies still needed to verify the positive results from the few initial clinical experiences available [29]. Resin Nano-Ceramic materials

Lava Ultimate CAD/CAM restorative is a highly cross-linked polymeric matrix contains a proprietary blend of three fillers: zirconia and silica nano particles agglomerated into clusters; individually bonded silica nano particles; and individually bonded zirconia nano particles. The material contains about 80% (by weight) of this filler blend. The result is a material that has different properties than feldspathic glass ceramic and composite materials. This new material offers higher flexural strength and fracture toughness, which results in long-term durability and maintains a polish well over the long term.

Lava Ultimate CAD/CAM restorative is easy to mill needs to polish for a few minutes to achieve an enamel-like luster. The material is available in eight shades, with both high and low translucency options, giving dentists the palette necessary to closely match the esthetics of the patient's dentition. It also allows for

easy adjustments. The material is suitable for use with crowns (including implant-supported crowns), onlays, inlays and veneers. The material can be used in challenging areas such as the second molar, and also blends very well esthetically alongside ceramic restorations [30]. Machining of Zirconia

CAD/CAM zirconia dental restorations can be produced according to two different machining techniques: soft machining of pre- sintered blanks or hard machining of fully sintered blanks Soft machining of pre-sintered blanks

Since its development, direct ceramic machining of pre-sintered 3 Y-TZP has become increasingly popular in dentistry and is now offered by a growing number of manufacturers. Briefly, the die or a wax pattern is scanned, an enlarged restoration is designed by computer software (CAD) and a pre-sintered ceramic blank is milled by computer aided machining then the restoration is then sintered at high temperature.

Typically the powder used in the fabrication of the blanks contains a binder that makes it suitable for pressing that will be eliminated during the pre-sintering step. The powders consist of spray-dried agglomerates (about 60^m in diameter) of much smaller crystals that are about 40nm in diameter. The blanks are manufactured by cold isostatic pressing. The mean pore size of the compacted

powder is very small in the order of 20-30nm with a very narrow pore size distribution [31].

The binder is eliminated during a pre-sintering heat treatment. This step has to be controlled carefully by manufacturers, as If the heating rate is too fast, the elimination of the binder and associated burn out products can lead to cracking of the blanks. Slow heating rates are therefore preferred. The pre-sintering temperature of the blanks affects the hardness and machinability. These two characteristics act in opposite directions: an adequate hardness is needed for the handling of the blanks but if the hardness is too high, it might be detrimental to the machinability. The temperature of the pre-sintering heat treatment also affects the roughness of the machined blank. Overall higher pre-sintering temperatures lead to rougher surfaces. The choice of a proper pre-sintering temperature is thus critical [31].

Several variations of the milling process exist for example; both contact scanners and non-contact scanners are available. Overall, non-contact scanners are characterized by a higher density of data points and a greater digitizing speed compared to contact scanners.

Sintering of the machined restorations has to be carefully controlled, typically by using specifically programmed furnaces. Shrinkage starts at 1000 °C and reaches about 25%. Sintering conditions are product-specific. Final sintering

temperatures between 1350 and 1550 °C with dwell times between 2 and 5hours lead to densities greater than 99% of the theoretical density. These variations in sintering conditions are likely to be due to the initial chemical composition of the 3YTZP powder. For example, small additions of alumina have been shown to act as a sintering aid, allowing the use of lower sintering temperatures and times. The minimum thickness for the copings is 0.5mm, below which warpage could occur. The restorations are furnace-cooled to a temperature below 200 °C to minimize residual stresses. As mentioned earlier, the sintering temperatures and times strongly influence the grain size [32]. Chevalier et al [11] also demonstrated that the amount of cubic phase in 3Y-TZP increases when the sintering temperature reaches 1500 °C with a sintering time of 5 h. The presence of larger cubic grains is detrimental to the resistance of the ceramic to low temperature aging.

Restorations can be colored after machining by immersion in solutions of various metal salts such as cerium, bismuth, iron or a combination thereof [33]. The color develops during the final sintering stage. The concentration of the solution strongly influences the final shade. Concentrations as low as 0.01mol% are sufficient to produce a satisfactory coloration. The final sintering temperature influences the color obtained. Careful respect of the manufacturer's instructions is therefore important. Coloration with various metal salts does not appear to affect the crystalline phases or mechanical properties of the final product. Alternatively,

colored zirconia can be obtained by small additions of various metal oxides to the starting powder [34]. The restorations are finally veneered with porcelains of matching coefficient of thermal expansion. The veneering porcelain is baked at nearly 900 °C, with a hold time of 1 min [9]. Representative systems utilizing soft machining of 3Y-TZP for dental restorations are Cercon® (Dentsply International), LavaTM (3MTM ESPETM), Procera® zirconia zirconia (Nobel BiocareTM), YZ cubes for Cerec InLab® (VidentTM) and IPS e.max® ZirCAD (Ivoclar Vivadent). Hard machining of 3Y-TZP and Mg-PSZ

Y-TZP blocks are prepared by presintering at temperatures below 1500 °C to reach a density of at least 95% of the theoretical density. The blocks are then processed by hot isostatic pressing at temperatures between 1400 and 1500 °C under high pressure in an inert gas atmosphere [35.36]. This latter treatment leads to a very high density in excess of 99% of the theoretical density.

The blocks can then be machined using a specially designed milling system. Due to the high hardness and low machinability of fully sintered Y-TZP, the milling system has to be particularly robust. A study by Blue et al. demonstrated that Y-TZP was significantly harder to machine than fully sintered alumina with lower material removal rates [37]. This was confirmed by other researchers who reported that coarse diamond burs were more efficient for material removal with Y-TZP, while machining with fine burs led to a more ductile type of damage [38,

39]. At least two systems, Denzir® (Cadesthetics AB) and DCZirkon® (DCS Dental AG) are available for hard machining of zirconia dental restorations.

Esthetic aspect of zirconia ceramics:

Zirconia framework is esthetically better accepted than metallic framework, but it remains too white and opaque. Colored zirconia framework was introduced to simulate the color of natural teeth. A traditional ceramic-engineering approach is to add coloring oxides to zirconia powders prior to blanks pressing. Another approach is color-doping of powder granules by co-precipitation, leading to homogeneous color distribution. The need for many shades makes both of these two approaches financially challenging. A popular means of coloring zirconia restorations is by infiltration of various metal salts at low concentrations [40]. However, this technique has some drawbacks, as a non uniform color due to the possible presence of porosity gradients [41]. Problems due to limited diffusion depth of coloring solutions have also been reported, leading to color variations from one area to another after grinding adjustments. Also, the sintering temperature should be carefully controlled, since it may influence oxidation state and therefore color [42].

Y- TZP blocks could be custom - colored where the technique involves infiltration of the machined restoration at the presintered stage being brought into a

highly porous state with special coloring solutions to produce work pieces of various shades [40].

These solutions could be applied through immersion of the restoration in coloring solution for a definite time and concentration, deposition of the solution by means of spraying process or through deposition of solutions by means of application instruments as brush or swab. After drying and at the initial stage of heating of the immersed porous presintered zirconia blocks, the acetic, chloric and nitric ions probably burn out or vaporized and disappear on the surface of the pores. The metal ions form an oxide layer on the surface of the pores of zirconia blocks. Theses metal oxides form a new product with zirconia via solid- solid reaction or remains as oxides at the boundaries around zirconia around zirconia grains in the final sintering stage (1350-1600°C). These techniques with infiltration with a coloring liquid although were thought to have theoretically more customization and characterization, few drawbacks as lake homogeneity and infiltration of the surface layers more than the bulk material had been reported [43]

The ability to control the shade of the zirconia core may eliminate the need to veneer the lingual and gingival aspects of the connectors in difficult situations like limited interocclusal distance and the required connector dimensions are minimally achieved. Also, the palatal aspect of anterior crowns and FPDs may be

fabricated of the core material only in cases like extensive vertical overlap and lack of space for lingual veneering porcelain [44.45].

Since the esthetic of value of a ceramic restoration is based on its ability to be in harmony with the adjacent natural teeth. Many authors considered the translucency as a key component for success of ceramic restorations [46,47]. Zirconia based ceramics have poor translucency due to their chemical nature, the amount of crystals, the particles' size, the pores and the sintered density. All of these factors determine the amount of light that is reflected, transmitted and absorbed and hence the optical properties of the ceramic materials [48]. Since Zirconia based ceramics (like Y-TZP) are polycrystalline materials , most of the light passing through it is intensely scattered and diffusely reflected leading to its opaque appearance[49].

Some studies suggested that microcrystals and full densification could enhance light transmission and optical properties of zirconia ceramic [50-52]. Jiang et al found that Y-TZP could gain nearly full density and 17-18% transmittance if finally sintered at 1450-1500°C [53].

Attempts trying to introduce translucent zirconia had been through consolidating nano-powder to full density with nano- crystals through the industrial sintering technique such as hot - isostatic pressing (HIP), millimeter wave, microwave and spark plasma sintering (SPS) [51-55]. Electric current assisted

powder consolidation techniques have became of the most successful methods for production of a fully dense ceramic with nano- size crystals[56]. Bonding of Veneering Material to Zirconia

Zirconia core is covered by veneering porcelains to realize the esthetic restoration. The veneering of the zirconia ceramic may be performed with ceramics using a layering technique, or a press technique, or combination of these techniques.

It seems that the veneering porcelains are mainly bonded to zirconia with mechanical interlocking and compressive stress due to the small difference between the zirconia and the veneering porcelain in thermal shrinkage by cooling after sintering. Also there was no evidence for chemical bonding between the zirconia and the veneering porcelains, SEM observation could not confirm the presence of reaction layer between the zirconia and the veneering porcelain. In addition to the difference in coefficient of thermal expansion many other factors could affect bonding between zirconia core and veneering material as poor wetting of the core with the veneering material, firing shrinkage of the veneering layer, phase transformation of zirconia crystals at the core - veneer interface due to thermal and / or mechanical stresses and inherent flaw formation during processing [57].

Coefficient of thermal expansion of veneering ceramic (8.8-10.0 x10 -6 per °C) which is slightly less than that of zirconia core (10.0-10.5 x 10-6 per °C), this difference leads to compression within the veneering mass and enhance bonding to zirconia core. Current studies demonstrated that sandblasting of the core material prior to veneer application did not enhance bonding, and even have negative effect due to alumina particles remain on the surface [58]. Bonding strengths between zirconia and veneering porcelains ranged between 26.5—31.6 MPa [50].

Chipping or lamination of the veneer material was recorded as one of the most common complications of zirconia restorations. Von steyern et al [59] reported that porcelain chip-off fractures causes the success rate of Dc zircon all-ceramic fixed partial denture to drop from 100 to 85% after 2 years. Also, Donovan reported that the fracture rate of the veneering ceramic has ranged from 8 to 50 percent at one to two years, while the reported rate of veneer fracture with metal ceramic restorations has been between 4 and 10 percent after 10 years [60]. The cause of this chipping is not known, and both core flexure and bond failure have been suggested. Another possible cause of chipping is the lack of uniform support of the veneering ceramic by the core. Bonfante et al [61] stated that their results suggest that the porcelain to core thickness ratio play a role on the fracture chip size in anatomically correct molar crowns, investigations concerning

alterations in core geometric design with thinner and more uniform veneer thickness are warranted.

One of the most frequent technical problem in all studies of zirconia reconstructions is chipping or cracking of the veneer ceramic In an investigation using a different prototype zirconia ceramic, veneer chipping was found in 15.2% of the cases after 35.1 months of follow-up [62]. In another investigation with the same zirconia ceramic, chipping was found in 13% after 37.2 months [63]. Also, chipping of the veneer was reported to occur in 15% of the cases after 2 years [64]. A fourth study reported chipping in 25% of cases after 31.2 months of observation [65]. Finally, a study reported the lowest chipping incidence, 6% after 37months. These results were attributed to low or moderate bond strength between zirconia frameworks and veneering ceramics and it can be concluded that various veneering ceramics available for zirconia possess insufficient mechanical properties [66].

The use of zirconia surface modification techniques to achieve strong bond between coping and veneering ceramic could improve the clinical failure rates observed. Application of a silicate intermediate layer, applied on the zirconia surface via a tribochemical approach or vapour deposition approach that could also enable conformal silicate surface modification without use of an aggressive physical process, which might result in damage to the coping surface. The

application of a liner, used to modify the colour of white zirconia for esthetics, has shown mixed results in bond strength when used on veneers [67].

Sandblasting is a popular means. It might be assumed that sandblasting would enhance the bond strength of cores to veneering porcelains by increasing surface roughness and providing undercuts [68,69]. The results of a recent study by Saleh [70] showed that sandblasting with 110 ^m Al2O3 particles effectively increased both surface roughness (Ra) of zirconia samples and shear bond strength between zirconia and the veneering porcelain compared with the untreated samples However, another study by kim et al reported that bond strength decreased by sandblasting [71].

Some studies have suggested that sandblasting induces transformation from the tetragonal to the monoclinic phase, thus affecting bond strength [72]. The coefficient of thermal expansion of monoclinic zirconia is 7.5*10-6/C, and that of tetragonal zirconia is 10.8*10-6/ C. Accordingly, an increase in the difference in the coefficient of thermal expansion between the TZP framework and the veneering ceramic leads to a decrease in bond strength [73]. Heat treatment of the TZP framework for 5 min at approximately 1,000 °C prior to commencement of the veneering process is recommended. Induction of the monoclinic zirconia phase by sandblasting is reversed by heat treatment [74], with a concomitant alteration in the coefficient of thermal expansion.

A recent evaluation of the shear bond strength of feldspathic porcelain to the Katana zirconia framework material concluded that shear bond strength between these materials was comparable to the bond between feldspathic porcelain and gold alloy and depends on the strength of the porcelain. The discrepancies in CTE between veneering porcelains and Katana zirconia significantly affected the shear bond strength of the porcelain zirconia system [75].

Based on concepts found in metal ceramic literature and known physical properties of zirconia, a zirconia collar may be used to support the veneering porcelain. Extending the zirconia collar to the interproximal area is useful to limit the amount of veneering porcelain that might fracture or chip [76]. Bonding and surface treatments

Several studies investigated different bonding methods as surface roughening, silica oating, silanization and the use of different bonding agents to zirconia ceramic.

Surface treatment methods causing miromechanical interlocking Sandblasting

Sandblasting is known to form surface roughness and irregularities and increase the surface area, wettability of the ceramic surface. Also it is used to clean the substrate surfaces, thus allowing resin cement to flow in to the surface. [77,78] Previous studies have demonstrated that sandblasting enhances the crown retention

regardless of the cement used. Although sandblasting improves bonding, its effect on mechanical properties is controversial some reports supposed mechanical weakening of zirconia [79] due to phase transformation which can cause fatigue in the material structure [80]. However, others reported that sandblasting strengthens the mechanical characteristics of zirconia [81-83]. Hot chemical etching

The action of hot etching solution is basically corrosion - controlled process where it selectively etches the zirconia ceramic surface resulting in rough surface and enhances the possibility of mechanical interlocking with resin cements. The etchant used modifies the grain boundaries by removing the less arranged high energy grain boundaries [84-86]. 1.2 Laser treatment

Absorption of laser energy by zirconia ceramic resulted in a smooth surface with irregular microcracks [87, 88]. Such surface treatment was investigated and the results were controversial, some investigations reported that laser treatment did not improve the mechanical interlocking with resin cement [87], while other reported that preparation of zirconia ceramic surfaces with CO2 and Er:YAG lasers significantly increased the bond strength, with the CO2 laser being superior to Er:YAG laser [89].

Nano- structured alumina coating

This method depends on the application of a nano-structured alumina coating with a high surface area and good wettability to create a micro-mechanical interlocking. The creation of the nano- structured layer is achieved by the hydrolysis of aluminum nitride to form lalamellar (AlOOH) onto the zirconia surface then a series of heat treatments were performed. The end results was a discontinuous nano- structured alumina coating (240^m thickness) which was found to significantly improve the bond strength to resin cements [90, 91]. Surface treatment methods causing chemical adhesion Silane coupling agents

Silane coupling agents are hydrid inorganic-organic synthetic compounds that are used to enhance the bonding of the resin composite to glass-based fillers and HF-etchable ceramics or silica coated metals and oxide ceramics [92]. Zirconia ceramics are not silica based and thus they present a physico-chemical challenge for reliable and durable resin bonding [93]. 2.2 Other coupling agents

A new primer (AZ Primer, Shofu, Kyoto, Japan) containing a phosphonic acid monomer, 6-MHPA (6-methacryloxyhexyl- phosphonoacetate), has been introduced for promoting bonding resin composite cements to alumina and zirconia ceramics. Kitayama et al. found that primers containing a phosphonic acid

monomer or a phosphate ester monomer, including 6-MHPA and MDP, were effective in promoting bonding of resin cements to zirconia ceramic. Without any primer, the resin cement containing MDP was found to be effective in bonding to zirconia ceramic [94]. 2.3 Resins and resin composites

Several studies suggesting the use of a phosphate monomer containing luting resin for cementation of zirconia restorations. It provides significantly higher retention of zirconia ceramic crowns than conventional luting cements, Tanaka et al. [95], Shahin and Kern [96], Kern and Wegner [97], de Oyague et al [98], reported that a phosphate monomer-containing luting system is recommended to bond zirconia and surface treatments are not necessary.

Manufacturers have produced new phosphate monomers designed not only to bond to zirconia. One of the recently developed phosphate monomers (RelyX Unicem) has a characteristic of self-etching phosphorylated methacrylates with two phosphate groups and at least two double bonded carbon atoms and this result in good bond strength to zirconia in addition to adequate crosslinking to the resin matrix [99]. Gas Fluorination

In an attempt to modify the zirconia to create a more reactive surface and hence improve the chemical bonding a new method in which an oxyfluoride

conversion layer was created on the surface of zirconia ceramic and thus increase its reactivity. This method was investigated by Piascik et al and found to be effective and they concluded that oxifluoration conversion layer formed was effective in bonding to silane coupling agent [100].

Surface treatment methods causing chemical bonding and micro- mechanical interlocking: Silica coating

There are several methods to coat zirconia ceramic surfaces with silica. The aims of this process are to clean the surface, create a highly retentive surface and the most important is to promote good bonding to resin cement through silane application.

There are many methods used for silica coating. A thermal silica-coating system (Silicoater MD system, Heraeus Kulzer, Wehrheim, Germany) was introduced since early 1980's. It involves sandblasting of the zirconia ceramic surface followed by silica-coating. Then, the surface was coated with silane which formed at increased temperature silica coating for the substrate in Silicoater MD apparatus [101].

Tribochemical silica coating utilizes the same principle with a specifically surface-modified alumina with SiO2 coating on the surface of the particles. The first tribochemical silica-coating system for dental use (Rocatec™ system, 3M

ESPE, Seefeld, Germany), introduced in 1989. It involves creating chemical bonds by applying kinetic energy in the form of sandblasting, without additional heat or light [102, 103].

Tribochemical silica-coating using CoJet™ at the dental office is a widely used now as a conditioning method for both ceramic and metal alloy. Literature recorded two possible disadvantages; possibility of subcritical crack propagation within zirconia in case of thin restorations [104-106], and the potential to remove significant amount of material which could affect their clinical adaptation during air abrasion step [107].

Other methods of silica coating include. Silicoater-technology based on pyrolysis of silanes is used in the PyrosilPen™-technology. They are based on a flame from mixture of butane gas and tetraethoxysilane (TEOS). Tetraethoxysilane decomposes in the flame to produce =Si-O-C= type species. Objects subjected for the flame over a short time (5 seconds), will be covered by a layer of these fragments which bond adhesively to the substrate surfaces. This surface has glass-like properties and be ready for resin bonding through silane application. So many researchers conducted in the last few years and proved the efficacy of the process of silica coating for Zirconia - resin bonding. A study by Ozcan and Vallittu [108] confirmed the effectiveness of silica coating and subsequent silane treatment on bonding to glass infiltrated zirconia. Also Sahafi et

al [109], confirmed the same finding by using tribochemical coating system on bonding to zirconia material. Same findings were recorded by Ernst et al. [110], Piwowarczyk et al [111], and Lüthy et al. [112]. Selective infiltration etching

Selective infiltration etching (SIE) is a novel surface treatment developed to transform zirconia surface into a retentive one by promoting the inter-grain nanoporosity. Such porosity will be infiltrated and interlocked by the resin cements. SIE utilizes a specific glass infiltration agent that is capable of diffusing between the grains and results in nano-inter-grain porosity. In a recent study, the effect of three different surface treatments on the bond strength was investigated: selective infiltration etching, particle abrasion, and coating with MDP monomer. The initial high bond strength was observed for the SIE surface treatment. Additionally, it was the only method capable of maintaining long-term bond strength values after storing in water for four weeks [113]. It was recorded by Kitayama et al. that the bond strength could be further improved with special zirconia primers [93].

Biological aspects of zirconia ceramics:

Biocompatibility of zirconium - based ceramics:

Biocompatibility has been defined as the ability of a material to perform with an appropriate host response in a specific application. Biocompatibility is one

of the most important advantages of zirconium based ceramics. Many In vitro tests were performed using cell cultures with cells such as fibroblasts, blood cells and osteoblast cells [114]. Garvie et al. implanted magnesia partially stabilised zirconia (Mg-PSZ) in paraspinal muscles of rabbits, which were examined at different intervals of time (1 week, 1 month, 3 months and 6 months). They reported no significant adverse soft tissue response to implants, and no degradation of the properties of Mg-PSZ occurred in contact with these living tissues. There was no change were detected in the surface phase content, surface roughness and strength of the implants [115].

Christel et al studied the effects of zirconia ceramic in vivo. Yttria stabilised zirconia and alumina cylinders were implanted into paraspinal muscles of rat. After 1 to 12 weeks after placement, no significant difference was observed between the two materials. They concluded that no cytotoxic effect for zirconia ceramic when tested both in vitro and in vivo [116]. Josset et al investigated human osteoblasts in culture with zirconia and alumina discs. They found that the osteoblasts showed good adhesion and spreading properties and the cells preserved their capacity to proliferate and differentiate into ostogenic pathways [117].

Scarano et al studied the bond response to zirconia implants inserted into tibia of white mature rabbits when the implants were retrieved after 4 weeks. They

reported the presence of newly formed bone and osteoblasts directly on zirconia implants with no inflammation observed [118].

Uo et al. studied cytotoxicity of dental ceramics cultured with human gingival fibroblast cells. They used Alamar Blue assay (the activity of the metabolic reaction of cells) and DNA quantification (number of cell growth and activity). They concluded that no cytotoxicity of various ceramics including Denzir (YTZP) [119].

Ichikawa et al. evaluated in vivo tissue reaction and stability of partially stabilized zirconia ceramic in subcutaneous implantation test. During the experimental period, zirconia ceramic was completely encapsulated by a thin fibrous connective tissue with (less than 80 ^m in thickness). No changes of weight and three-point bending strength of the zirconia samples were detected after 12 months of implantation. The result of this study suggested that zirconia ceramic is biocompatible and no degradation of zirconia ceramic occurred [120].

Researchers reported that zirconia is not able to generate mutations of the cellular genome [121], in particular, mutant fibroblasts found on ZrO2 were fewer than those obtained with the lowest possible oncogenic dose compatible with survival of the cells [122]. Moreover, zirconium oxide creates less flogistic reaction in tissue than other restorative materials such as titanium [123]. This result was also confirmed by a study about peri-implant soft tissue around zirconia healing caps in

compared with that around titanium healing caps. Inflammatory infiltrate, microvessel density, and vascular endothelial growth factor expression were found to be higher around the titanium caps than around the ZrO2 ones [124].

Bacterial adhesion, which is an important aspect in order to maintain zirconia restorations without marginal infiltrations or periodontal alterations, proved to be satisfactorily with zirconia restorations. Scarano et al. found a degree of coverage by bacteria of 12.1% on zirconia as compared to 19.3% on titanium. [125]. Rimondini et al. confirmed these results with an in vivo study, in which Y-TZP accumulated fewer bacteria than Titanium in terms of the total number of bacteria and presence of potential putative pathogens such as rods [126].

Although Zirconia materials showed excellent tissue reaction, but in the same time this material has no direct bone bonding properties or osteoconduction behavior, except it shows a morphological fixation with the surrounding tissues [127,128]. Strategies for modifying the nature of the material surfaces may affect cell-material interactions, which in turn influence tissue development. The results obtained by Wu et al showed that the wettability of ZrO2 were noticeably improved by oxygen plasma treatment that was effective for retaining a stable surface hydrophilicity. The improved wettability did actively promote the attachment, proliferation and differentiation of human osteoblast-like cells [129]. Radioactivity of zirconium based ceramics:

Zirconia powder contains small amounts of radionuclides from the

OOfy Q-1Q___

uranium, radium ( Ra) and thorium ( Th) actinide series. The separation of these elements is difficult and costly. Two types of radiation are correlated with zirconia, alpha and gamma. Significant amounts of alpha radiation have been observed in zirconia based ceramics used in the manufacture of implants and due to their high ionization; the alpha particles destroy cells of hard and soft tissues but the literature suggests that the radiation level is not worrisome in zirconia [130].

Great concerns were raised in the early 1990's about the use of zirconia ceramics for medical and dental applications because of these radioactive impurities and their effect on human body. However, after purifying procedures, zirconia powders with low radioactivity (< 100 Gyh-1) can be achieved [131]. These values are below the European radiation limits accepted for human body external exposure or for local internal exposure of organs and tissues and so near to those of alumina ceramics and Co-Cr alloys for dental and medical uses. After purifying of zirconia powders, uranium concentration in zirconia powders ranges between 0.001 and 0.007 Bq/g, (max: 1.0 Bq/g according to ISO 6872). Therefore, the radiation levels of the commercially available zirconia powders conform to the recommendations of the International Commission for Radiation Protection (ICRP) and are generally

found lower than the normal ambient radioactivity induced by natural radiations [131].

Moreover, The mass activity of the current zirconia powders is much lower than the mass activity of the human bone (20-50 Bq/kg), which roughly corresponds to a radionuclide content of 1 ppm (0.0001%) or less which is quite below the values recommended by the International Standards Organization (ISO 13356) and the American Society for Testing and Materials Standards (F1873). They recommended that the mass activity of zirconia powders used for surgical grade ceramics should be lower than 200 Bq/kg [132]. Conclusions:

1. Dental applications of zirconia ceramics seem to gain a well- established place in clinical dentistry today, due to its exceptional physical and biological properties and the great enhancement in CAD-CAM technology.

2. A reliable choice of surface treatment and luting cement for cementation of zirconia contributes toward preventing leakage. It is recommended to use resin cements because they have been shown to produce higher bonding to both the zirconia surface and dentin when compared to conventional cements.

3. Because of the rapid development of both materials and processing technologies, application of zirconia-based FPDs can be seen as promising alternative. However, dentists, researchers and dental technicians must collaborate to overcome the esthetic limitation of zirconia materials.

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