Mechanical and Thermal Cycling Effects on the Flexural Strength of Glass Ceramics Fused to Titanium

Dental Materials Journal 27(1)：7－15, 2008

Mechanical and Thermal Cycling Effects on the Flexural Strength of Glass Ceramics Fused to Titanium

Vanessa VÁSQUEZ1, Mutlu ÖZCAN2, Renato NISHIOKA1, Rodrigo SOUZA1, Alfredo MESQUITA1 and Carlos PAVANELLI1

1São Jose dos Campos Dental School, Department of Dental Materials and Prosthodontics, São Paulo State University, São Jose dos Campos, Brazil2University Medical Center Groningen, University of Groningen, Department of Dentistry and Dental Hygiene, Clinical Dental Biomaterials, Antonius Deusinglaan 1, 9713 AV, Groningen, The NetherlandsCorresponding author, Mutlu Özcan; E-mail: [email protected]

Received April 4, 2007/Accepted July 20, 2007

　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　 This study evaluated the effects of mechanical and thermal cycling on the flexural strength (ISO 9693) of three brands of ceramics fused to commercially pure titanium (cpTi). Metallic frameworks of 25×3×0.5 mm dimensions (N＝84) were cast in cpTi, followed by 150-μm aluminum oxide airborne particle abrasion at a designated area of the frameworks (8×3 mm). Bonder and opaque ceramic were applied on the frameworks, and then the corresponding ceramic (Triceram, Super Porce-lain Ti-22, Vita Titankeramik) was fired onto them (thickness: 1 mm). Half of the specimens from each ceramic-metal com-bination were randomly tested without aging (only water storage at 37ºC for 24 hours), while the other half were mechani-cally loaded (20,000 cycles under 10 N load, immersion in distilled water at 37℃) and thermocycled (3,000 cycles, between 5－55℃, dwell time of 13 seconds). After the flexural strength test, failure types were noted. Mechanical and thermal cycling decreased the mean flexural strength values significantly (p<0.05) for all the three ceramic-cpTi combinations tested when compared to the control group. In all the three groups, failure type was exclusively adhesive at the opaque ceramic-cpTi interfacial zone with no presence of ceramic on the substrate surface except for a visible oxide layer.

Keywords: Flexural strength, Aging, Titanium　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　　

INTRODUCTION

Although there is a growing trend for metal-free res-torations in the dental profession, failures associated with such materials are also being reported ― indi-cating that there is still place for metal-ceramic fixed partial dentures (FPD). Furthermore, extensive oral rehabilitation could be achieved with metal-ceramic FPDs only. However, the high cost of noble alloys and the potential biological hazards of base metal alloys pose some concerns for practical application. To circumvent these concerns, commercially pure titanium (cpTi) and some of its alloys were intro-duced for the construction of dental prostheses1).　　According to the American Society for Test-ing and Materials (ASTM), cpTi is available in four different grades (grade I to IV). It is based on the incorporation of small amounts of oxygen, nitrogen, hydrogen, iron, and carbon during purification pro-cedures, whereby each grade has different physi-cal and mechanical properties. Grades I and II are the most commonly used types in the production of metal-ceramic FPDs1). In implants and implant-sup-ported FPDs, cpTi and its alloys exhibit remarkable advantages due to their excellent biocompatibility, corrosion resistance, high strength, and low modulus of elasticity1-4).　　However, cpTi has a high melting point and it is active at high temperatures. For these reasons, it may become fragile if it reacts with atmospheric

oxygen during casting when conventional casting and investing methods are used5). Furthermore, tempera-tures above 800℃ increase the oxygen-rich titanium oxide layer on the surface, so called α-case, which impairs the mechanical compatibility of the titanium-ceramic system6-9). It should also be mentioned that cpTi has a low thermal expansion coefficient that makes it a difficult substrate for ceramics to bond onto.　　To the end of better compatibility with cpTi, ceramics to meet this specific purpose have been developed1). These ceramics, so called low-fusing ceramics (LFC), melt at temperatures lower than 800℃. Moreover, they have thermal expansion coef-ficients close to that of titanium, thereby reducing the thermo-mechanical stresses at the interface and allowing satisfactory bonding of the two elements10). By virtue of the advances in dental technology, the α-case layer can also be controlled almost com-pletely through improved casting techniques. This is done by induction in an inert atmosphere of argon or helium gas and with the use of refractory invest-ments that contain oxides such as magnesium, yttrium, and zirconium11-13).　　The success of metal-ceramic FPDs hinges on a durable adhesion between ceramic and the metal substructure9,14,15). In today’s dentistry, cpTi compat-ibility with ceramics, refractory materials, and spe-cial casting systems are not widely studied. These studies should indeed be undertaken with a view to

Ceramic-titanium flexural strength8

improving the success rate of metal-ceramic FPDs1). The cpTi-ceramic adhesion could be tested in sev-eral ways. These tests can be classified according to the nature of the applied stress, such as shear, ten-sile, flexural strength, or torsion test. However, to date, there is still no consensus as to which method is more valid for clinical applications16). In addition, these methods often do not involve the fatigue com-ponent of mechanical tests, thereby rendering the results unrealistically optimistic to a certain extent.　　To better predict the clinical behaviors of dif-ferent materials and material combinations, their mechanical fatigue tests should be carried out in a wet environment17,18). In this connection, thermo-cycling that is based on temperature alternations induces repeated stress to the metal-ceramic inter-face18,19). During thermocycling, differences between the thermal expansion coefficients of the two com-ponents could affect adhesive strength. However, a combination of both mechanical and thermal cycling could be considered a more aggressive way of aging the ceramic-metal interface.　　Undisputedly, cpTi and its alloys have many advantages and favorable properties over the other dental alloys. However, controversies exist in the lit-erature regarding the adhesion of ceramics to these alloys7,20-22). Therefore, the objectives of this study were twofold: (1) to evaluate the effects of mechanical and thermal cycling on the flexural strength of three brands of ceramics fused to cpTi; and (2) to asses the type of failure at the ceramic-alloy interface.

MATERIALS AND METHODS

Materials usedThree brands of ceramics ― namely, Triceram, Super Porcelain Ti-22, and Vita Titankeramik ― were used in combination with cpTi in this study. Brand names, manufacturers, and batch numbers of the ceramic types and titanium are presented in Table 1.

Fabrication of metallic frameworksRectangular acrylic templates (25×3×0.7 mm) were

used for the fabrication of frameworks. A wax sprue (Horus, Herpo Produtos Dentários Ltd., São Paulo, Brazil) was attached perpendicular at one end of the template and connected to a central wax rod of 5 mm diameter (Wax Wire for Casting Sprues, Dentaurum, Pforzheim, Germany). The assembly was mounted in a silicone ring and poured with investment material (Rematitan® Ultra, Dentaurum JP Winkelstroeter KG, Pforzheim, Germany) that was mixed at a ratio of 100 g of powder to 14 ml of liquid. After the investment material set, the silicone ring and sprue former were separated from the investment mold. Metallic frameworks were cast in cpTi (N＝84) in an electrical induction furnace (Rematitan® Autocast, Dentaurum) under argon gas. Elimination of sprues and separation of metallic strips were performed with the aid of carbide disks at low speed.　　After removal from the investment mold, mar-gins of the frameworks were trimmed to the final dimensions of 25×3×0.5 mm. Surfaces of the speci-mens that would receive the ceramic layer were treated with airborne particle abrasion with 150-μm aluminum oxide (Korox, Bego, Bremen, Germany). This was done at an angle of 45°for 10 seconds, and from a distance of approximately 2 cm under 2 bar pressure. Frameworks were then ultrasonicallycleaned in isopropyl alcohol (Vitasonic II, Vita Zahnfabrik, Bad Säckingen, Germany) for five minutes and allowed to dry at room temperature.

Application of ceramic layerAn area of 8×3 mm was marked on the cpTi frame-works with a graphite pencil. Then, the bonder of each ceramic was applied in a thin layer with a brush. After firing, opaque ceramic paste ― con-sisting of opaque ceramic powder mixed and homog-enized with liquid in a container connected to a dis-penser ― was applied on the bonder. Thickness of the ceramic layer corresponding to dentin ceramic (1 mm) was standardized by positioning the frame-works in a metallic template. After removal from the assembly, ceramic was fired. Due to shrinkage, a second layer was applied and then the specimens

Brand name Ceramic Type Manufacturer Batch number

Triceram Low-fusing ceramic Dentaurum, Ispringen, Germany 003

Super Porcelain Ti-22 Low-fusing ceramic Noritake, Nagoya, Japan 60506

Vita Titankeramik Low-fusing ceramic Vita Zahnfabrik, Bad Säckingen, Germany

2370

Commercially pure titanium (Tritan) Dentaurum J.P. Winkelstroeter KG, Pforzheim, Germany

098

Table 1 Brand names, indications, compositions, and manufacturers of dental ceramics and titanium used in this study

VÁSQUEZ et al. 9

were submitted for final glaze firing (Table 2).　　Twenty-eight specimens for each ceramic-cpTi combination were randomly divided into two sub-groups: mechanical and thermal cycling versus only stored in distilled water for 24 hours at 37℃ (control group) prior to flexural strength test.

Mechanical and thermal cyclingMechanical cycling of the specimens was carried out in a mechanical cycling machine (custom made, Paulista State University, Dental School, UNESP, Sao Jose dos Campos, Brazil). The latter was devel-oped to simulate the mechanical forces generated during the chewing cycle. The device employed for this test was composed of two bases, 2 cm apart from each other, on which cylinders (radius: 1.0 mm) were placed to allow specimens to be positioned parallel to the ground and perpendicular to the axial load. An upper rod with a 1 mm diameter tip was fixed on a plier to apply a 10 N load for 20,000 times at a fre-quency of 1 cycle per second. The device for testing was placed on the machine base that contained a thermostat to allow testing in an aqueous medium at a constant temperature of 37℃.　　Subsequently, the specimens were thermocycled for 3,000 cycles between 4℃ and 55℃ in deionized

water (Nova Etica, São Paulo, Brazil). Dwelling time at each temperature was 10 seconds, and the trans-fer time from one bath to the other was five seconds.　　The process of subjecting the specimens to 20,000 times of mechanical cycling followed by 3,000 times of thermal cycling is regarded as the“aging” procedure.

Flexural strength testFlexural tests were performed in a universal test-ing machine (Instron 4301, Instron Corp., Norwood, MA, USA). The load was applied at a constant speed of 1.5 mm per minute until fracture occurred (Figs. 1(a)－(d)). The formula according to the guidelines of ISO 969323) was adopted for the calculation of data obtained from the flexural strength test. The load that led to initial separation of materials was obtained in kilogram force (kgf). It was converted to Newton (N) for the calculation of flexural strength according to the following equation:

where P is the maximum load upon fracture (N), I the distance between two supports (mm), and b the

Ceramics Starting Temperature (ºC)

Drying Time (min)

Final Temperature (℃)

Temperature rate of increase (℃/min)

Holding time (min)

Triceram

Bonder 500 4 795 65 1

Opaque 500 4 795 65 1

First dentine layer 500 6 755 40 1

Second dentine layer 500 4 755 40 1

Super Porcelain Ti-22

Oxidation 500 3 800 50 3

Bonder 500 5 800 50 1

Opaque 500 5 780 50 1

First dentine layer 500 7 760 40 1

Second dentine layer 500 7 760 40 1

Vita Titankeramik

Bonder 400 6 800 60 1

Opaque 400 2 790 110 1

First dentine layer 400 6 770 50 1

Second dentine layer 400 6 770 50 1

Table 2 Firing procedures of the dental ceramics tested

Ceramic-titanium flexural strength10

width (mm) and d the thickness (mm) of the speci-men.　　Specimens were analyzed under a stereomi-croscope (Stemi 2000-C, Carl Zeiss, Gottingen, Ger-many). Images were digitally recorded with a cam-era (Cybershot, Model DSC-S85, Sony, Tokyo, Japan), which was connected to the microscope, for metal surface characterization and assessment of failure modes. In this study, failure types were classified as follows: (a) adhesive failure along the interfacial region between opaque ceramic and ceramic-metal interaction zone; (b) inside the interaction zone; and (c) cohesive failure along the interfacial region between metal and the interaction zone9).

Statistical analysisStatistical analysis was performed using STATIS-TICA for Windows (version 5.5, StatSoft Inc., Tulsa, OK, USA) and Statistix for Windows (version 8.0, Analytical Software Inc., Tallahassee, FL, USA). The means of each group were analyzed by two-way

analysis of variance (ANOVA), with flexural strength as the dependent variable and the ceramic-metal combination and aging condition as the independent factors. P values less than 0.05 were considered to be statistically significant in all the tests. Multiple comparisons were made by Tukey’s adjustment.

RESULTS

Table 3 presents the two-way analysis of variance (ANOVA) results for the experimental conditions. Interaction between ceramic type and aging factor was not statistically significant (p＝0.168) (ANOVA, Tukey’s test). In Table 4, it was shown that mechan-ical and thermal cycling decreased the mean flex-ural strength values significantly (p<0.05) for all the three ceramic-cpTi combinations tested (27.4±4.1－28.5±5.4 MPa) when compared to the control group. In the latter group, tests were performed after 24-hour water storage at 37℃ (32.5±4.3－38.5±2.3 MPa).

Fig. 1 (a) Final shape and dimensions of ceramic-alloy specimen; (b) cross-section dimensions of specimen according to ISO 9693; (c) application of axial force; (d) separation of ceramic from titanium surface.

VÁSQUEZ et al. 11

　　Results of Tukey’s multiple comparison test established that only Vita Titankeramik-cpTi system showed a significantly higher value than those of other ceramic-cpTi combinations in the control group (p<0.05) (Fig. 2).　　Stereomicroscope images at ×25 magnifica-tion showed exclusively adhesive failure mode at the opaque ceramic-cpTi interfacial zone in all the exper-imental groups (Triceram-cpTi, Super Porcelain Ti 22-cpTi, and Vita Titankeramik-cpTi metal-ceramic

systems). There was no presence of ceramic on the substrate surface, except for a visible layer of dark titanium oxide. Representative images of original cpTi-ceramic substrate and cpTi surfaces after flex-ural strength test are illustrated in Figs. 3(a)－(g).

Fig. 2 Mean flexural strength values according to the experimental conditions established by the variables of ceramic type and aging (20,000 times of mechanical cycling followed by 3,000 times of thermal cycling).

Experimental Groups

Mechanical- and thermal-cycling Mean (SD)

Without WithTriceram-cpTi (Group 1)

32.5±4.3a,b,c 28.4±5.3b,c 30.5±5.1

Super Porcelain Ti-22-cpTi (Group 2)

35.4±7.6a,b 28.5±5.4b,c 32±7.3

Vita Titankeramik-cpTi (Group 3)

38.5±2.3a 27.4±4.1c 32.8±6.8

Mean (SD) 35.5±5.5 35.5±5.5* Means followed by equal letters do not differ statistically

Table 4 Mean (±standard deviation) flexural strength values (MPa) of ceramic-cpTi combinations with and without mechanical and thermal cycling. Same superscript letters indicate no significant differences (Tukey’s test, α＝0.05)

Effect DF SS MS F P

Cycling 1 594.00 594.00 22.78 0.001*Groups (metal-ceramic)

2 38.21 19.11 0.73 0.488

Interaction 2 97.76 48.88 1.87 0.168

Residue Total

36 938.74 26.08

41 1668.72

* Statistically significant difference at the level of 5%.

Table 3 Results of two-way analysis of variance (ANOVA) for ceramic type, cycling fatigue conditions, and their interaction according to flexural strength data (*: p<0.05)

Ceramic-titanium flexural strength12

Fig. 3 (a) Representative image of ceramic specimen fired onto cpTi and optical microscopic images of specimens (×25) after flexural strength test: (b)－(c) Triceram-cpTi; (d)－(e) Super Porcelain Ti-22-cpTi; (f)－(g) Vita Titankeramik-cpTi. Clear appearance of cpTi metal surface is seen as well as blur appearance of oxide layer on the ceramic surface. Metal and oxide surfaces are indicated by an arrow for each ceramic-cpTi combination

VÁSQUEZ et al. 13

DISCUSSION

The clinical performance of metal-ceramic FPDs is usually estimated by mechanical strength tests, which assess the adhesion between a ceramic and the metal substrate. The nature of metal-ceramic bond-ing has been studied extensively, and it is funda-mentally based on three mechanisms: micromechani-cal retention, compressive adaptation, and chemical bonding9,15,24).　　In dental materials research, flexural strength test is considered as the most appropriate method to measure bond strength between two materi-als4,10,14,18,19,21,25-28). Some authors would recommend three-point bending test29,30), while others would advocate four-point or biaxial flexure strength test to evaluate metal-ceramic adhesion31). According to ISO 969323), flexural strength test should be preferred over the other methods because the stresses it simu-lates closely represent the stresses that occur in den-tal prostheses with multiples elements.　　The α-case layer that formed on the surface of melted cpTi alloys provided good chemical stability. However, at the same time, this intermediate oxide layer could impair the chemical reactions between cpTi and the bonder ceramic components, thereby weakening the cpTi-ceramic bonding. It is note-worthy that this oxide layer is often not strongly adhered to the metal surface and that it is porous, brittle, and incompatible with ceramics. It was therefore little wonder that the oxide layer itself or the interface between the oxide layer and the alloy has been shown to be responsible for metal-ceramic fractures3,6,14,28). In a bid to circumvent the incompat-ibility problem between the α-case layer and ceram-ics, a ceramic bonder corresponding to each ceramic material was thus used in this study.　　Besides, composition of the investment mate-rial and casting procedure also influence the forma-tion of α-case layer. For this reason, some authors advocated the use of Mg-based investment materi-als32), while others suggested Zr-modified Mg-based investment materials33). Addition of these elements increases the thermal expansion coefficient of the investment material and limits its interfacial reactiv-ity with cpTi at high temperatures. In the present study, the investment material used was in compli-ance with the manufacturer’s recommendations.　　The pressure at which liquid titanium was injected into the investment32), and the temperature at which this procedure was performed, could also influence the consequential adhesion of ceramics to cpTi22). Within this context, ceramics have been spe-cifically developed to make their firing onto cpTi pos-sible at temperatures below 800℃, so as to prevent excessive formation of oxides on the metallic frame-work surface2). The ceramic systems employed in

this study presented low firing temperatures for each layer ― between 750 and 800℃. Lautenschlager and Monaghan2), Wang and Fenton1), and Esquivel et al.10) proposed improved adhesion between ceram-ics and cpTi with so-called ultra-low fusing ceramics, whereby the firing temperatures range between 650 and 850℃9). These ceramics decrease the transfor-mation phase of cpTi, resulting in a stable oxide layer and good adhesion of the ceramics. Failure analysis also indicated a visible dark zone of oxide layer.　　Differences in thermal expansion coefficient between the metal and ceramic can produce resid-ual stresses along the interface, which can result in debonding or fracture of the ceramic9,21,22). Thermal expansion coefficients of the ceramics used in this research varied between 8.4 and 8.9×10－6K－1, being very close to each other. Thus, non-significant differ-ences among the ceramics tested could be explained on this ground. In the present study, results of the control group without mechanical and thermal aging (32.5－38.5 MPa) corroborated with the findings of Yilmaz and Dincer14) and White et al.17) ― where low-fusing ceramics were also used. However, between the studies of Pröbster et al.19) and Yilmaz and Dincer14), the flexural strength values showed great variations ranging from 14 MPa to 37 MPa. None-theless, results of the current study ― both the con-trol group without aging and even after aging ― were higher than the recommended minimum value of 25 MPa as established by DIN 13.92734).　　In terms of specimen dimensions, Yoda et al.4) suggested that the specifications of ISO 969323) did not entirely represent those of metal-ceramic dental restorative systems and Troia et al.28) attributed this to non-uniformity of ceramic thickness. For this rea-son, ceramic thickness was standardized at 1 mm in this study.　　The effects of mechanical and thermal cycling simulate ― to a certain extent ― the effects of the clinical service which dental materials are subjected to in the oral environment. In dental materials research, most in vitro experiments are performed using static mechanical tests that do not address the aggressive oral environment. It is known that the oral environment is able to induce physicochemi-cal alterations in dental materials. Temperature changes provide conditions for occurrence of deg-radation in an aqueous environment35). They also encourage mechanical fatigue of the materials them-selves or their interfaces, which is triggered by the repeated chewing action36,37). As for the water stor-age of ceramic materials, it decreases the latter’s mechanical properties38). The reduction may be related to the solubility of different oxides. This process may be higher in ceramics designed for use with cpTi because of the presence of alkaline metallic oxides35) ― which was also evident from the failure

Ceramic-titanium flexural strength14

analysis of this study.　　Previous studies which investigated the adhesion between metal and laboratory resin or between metal and ceramics disclosed reduced adhesion after ther-mal cycling18,19,39-41). In the study of Tróia Jr et al.28), the same form of thermal variation and metallic sub-strate (cpTi) as the present study were investigated. However, it was found that thermal cycling did not exert any influence on adhesion of ceramic with this metal. Nonetheless, this factor became more sig-nificant and pronounced in more drastic conditions as those employed in the study of Shimoe et al.42), where extended thermal cycling (100,000 cycles) was performed. As a result, a 30％ reduction in mean bond strength was observed. In this study, the chief focus was on the combined effects of mechanical and thermal cycling. On this ground, extended thermal cycling was not performed.　　According to Scherrer et al.36), all materials and their combinations should be subjected to fatigue conditioning before mechanical testing is performed. Mechanical fatigue is a helpful means to predict the clinical behavior of dental restorative systems in order to avoid catastrophic failures in vivo37). Some factors inherent to mechanical cycling may also influence the outcome. In this study, the applied load was 10 N and the number of cycles was 20,000 times at a frequency of 1 Hz as described by Itinochi et al.43). As for thermal cycling, the number of cycles adopted and the temperature alterations employed were based on previous reports18,19,28). Thermal cycling induces repeated stress at the metal-ceramic interface, thereby weakening the bond between the two components28,44,45). In this connection, Tróia et al.28) suggested that extended immersion times in each bath might produce higher tension at the metal-ceramic interface. On this note, Pröbster et al.19) and Leibrock et al.45) suggested that 6,000 thermal cycles would correspond to five years in physiological condi-tions.　　In this study, reduction in mean flexural strength was evident after aging. Although previous studies that investigated ceramic-cpTi bond strength did not incorporate the aging conditions19,28), other studies sought to achieve the same research objective by modifying fatigue variables such as load magni-tude and type of load (such as repetitive or dynamic load)18). Therefore, with a view to making study results comparable, there seems to be a need to stan-dardize the aging parameters.　　In clinical situations, many metal-ceramic FPD failures are in the form of ceramic fractures with metal exposure. While the search is still on for a better ceramic-compatible metal/alloy with improved clinical performance, clinicians should in the mean-while bear in mind that ceramic-metal bonding is susceptible to mechanical and thermal fatigue.

CONCLUSIONS

Mechanical cycling for 20,000 times followed by ther-mal cycling for 3,000 times decreased the mean flex-ural strength values significantly for all the three ceramic-cpTi combinations tested when compared to the control group ― where tests were performed after mere 24 hours’ water storage at 37℃. All tested ceramic-cpTi combinations showed adhesive fail-ure with a visible oxide layer on the cpTi substrate, indicating that the weakest link was still between ceramic and the titanium oxide assembly.

ACKNOWLEDGEMENTS

This investigation was financially supported by a grant, Grant No. 01166/04, from FUNDUNESP (Fundação para Desenvolvimento da Universidade Estadual Paulista).

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