factors affecting the fracture strength and fatigue ... · ii factors affecting the fracture...

Factors Affecting the Fracture Strength and

Fatigue Resistance of Molar Crowns Produced

Using a Zirconia-based System

by

Mohammed Hani Zahran

A thesis submitted in conformity with the requirements

for the degree of Doctor of Philosophy

Graduate Department of Dentistry

University of Toronto

© Copyright by Mohammed Hani Zahran (2013)

ii

Factors Affecting the Fracture Strength and Fatigue Resistance of Molar

Crowns Produced Using a Zirconia-based System

Mohammed Hani Zahran

Doctor of Philosophy

Graduate Department of Dentistry

University of Toronto

2013

Zirconium-oxide ceramic was introduced as a potential restorative material for

posterior teeth due to its superior mechanical properties. However, zirconia is opaque and

needs to be masked with an esthetic veneer. Short-term clinical studies showed a high

rate of veneer chipping/delamination for zirconia-based posterior crowns. This may be

attributed to either inferior strength of the veneering porcelain or weakness in the

zirconia-veneer bond. The aims of this in-vitro project were to analyze some of the

factors that can influence the performance of zirconia-based molar crowns and to

characterize the zirconia-veneer interface.

The project consisted of four parts. The first part aimed to evaluate the effect of total

thickness, core thickness and veneering technique on fracture mode and strength of

zirconia-porcelain specimens of simple geometry. Results indicated that all three

variables significantly affected mean fracture load, while only two (total thickness and

core thickness) affected fracture mode.

The second part aimed to assess the effect of core design and thickness on fatigue

resistance, fracture strength and fracture mode/extent of zirconia-based molar crowns.

Results indicated that core thickness, but not core design, had a significant effect on

iii

fracture strength. The anatomically-shaped core design reduced the size of the veneering

porcelain fracture.

In the third part, the effect of core design and two different CAD-CAM-produced

veneering materials on fatigue resistance, fracture load and fracture mode of zirconia-

based molar crowns was examined. Results indicated that the veneering material, but not

core design, had a significant effect on fracture load and mode.

The fourth part aimed to characterize the zirconia-porcelain interface using modern

surface analysis techniques (Time of Flight Secondary Ion Mass Spectrometry (ToF-

SIMS) and X-ray Photoelectron Spectroscopy (XPS)). Results indicated chemical

changes of zirconium along the interface.

In conclusion, improving the strength of the veneering layer by using pressed

veneering porcelain can enhance performance of posterior zirconia-based crowns in

comparison to manual build-up technique. Furthermore, fusing a CAD-CAM-produced

veneering layer to the underlying zirconia core is a promising technique that can augment

veneering porcelain strength and enhance zirconia-porcelain bond. Modifying core

design, crown thickness and core thickness can improve performance of zirconia-based

molar crowns.

iv

Dedication

v

Acknowledgement

Praise and gratitude to ALLAH, Almighty, without whose gracious help it would have

been impossible to accomplish this work.

First and foremost, I would like to express my gratitude to my supervisor, Prof. Omar

El-Mowafy, whose expertise, understanding, time and patience, added considerably to

my graduate experience. I appreciate his support, advice, guidance and mentorship

throughout my master and doctorate journey, which lasted more than 8 years.

Special thanks to my co-supervisor, Prof. Asbjorn Jokstad, for his expertise,

understanding, patience, professionalism and mentorship during my clinical training at

the Faculty. His insight and guidance contributed significantly to this project.

Also, I wish to thank my other thesis advisory committee members, Dr. Laura Tam

and Dr. Amin Rizkalla for the time they provided throughout this project. Dr. Tam’s

valuable and thorough comments and suggestions helped significantly in improving the

quality of this finished product. Without Dr. Rizkalla’s expertise in the field of

engineering and his appreciated effort in performing some of the mechanical testing at

his lab, this project would not be possible.

Specials thanks also go to my family for the support they provided through my entire

life, and very special thanks go to my wife, love and best friend, Dania, and my sons,

Abdullah, Ibrahim and Yousuf. Without their love, encouragement, on-going support and

sacrifices throughout this process, I would not have finished my graduate program. A

special welcome goes to the new member of our family, Abdulelah, who was hidding for

the last 9 months and joined us just after defending my dissertation.

In conclusion, I recognize that this research would not have been possible without the

financial support of the King Abdulaziz University, Saudi Arabia. This support is duly

acknowledged.

vi

Table of Contents

Chapter 1: Introduction and Literature Review ...................................................................... 1

1.1. Introduction ...................................................................................................................................................... 2

1.2. Dental Ceramics Classification .................................................................................................................. 3

1.3. Indications for All-Ceramic Restorations ............................................................................................. 4

1.4. Fracture of All-Ceramics Materials and Restorations ..................................................................... 4

1.5. Fatigue Process in Ceramics Materials and Restorations .............................................................. 5

1.5.1. Factors Affecting the Fracture and Fatigue of Ceramic Restorations ....................................... 7

1.5.2. Improving the Strength of Ceramic Material .................................................................................... 12

1.5.3. Laboratory Testing of the Performance of the Ceramic Materials and Restorations ...... 13

1.5.4. Methods for Testing the Effect of Fatigue on Ceramic Materials and Restorations ......... 14

1.5.5. Simulating the Clinical Situation in the Fatigue Testing .............................................................. 18

1.6. Survival Rates of All-Ceramic Crowns .................................................................................................. 19

1.7. Zirconium Oxide Ceramics ........................................................................................................................ 20

Chapter 2: Rationale and Objectives .................................................................................... 27

Chapter 3: Manuscript 1 ....................................................................................................... 29

Effect of total thickness, core thickness and veneering porcelain application technique on

fracture of zirconia/porcelain combinations ............................................................................................ 30


Effect of core thickness and design on fracture strength and fatigue resistance of

zirconia-based crowns ........................................................................................................................................ 59


Effect of core design and veneering material on fracture strength and fatigue resistance

of zirconia molar crowns ................................................................................................................................... 97

Chapter 6: Manuscript 4 ..................................................................................................... 129

Characterization of the interface between zirconia and veneering porcelain .......................... 130

vii

Chapter 7: Summary and Clinical Relevance ..................................................................... 155

Appendices ......................................................................................................................... 159

Reference ............................................................................................................................ 197

viii

List of Tables

(Table 3.1) Composition and some properties of materials used for core and veneering

porcelain. ...................................................................................................................... 44

(Table 3.2) Firing program for materials used in the study. ................................................. 45

(Table 3.3) Experimental groups according to core thickness (mm), total thickness (mm)

and veneering porcelain. ............................................................................................... 46

(Table 3.4) Fracture load (N) data for all groups. ................................................................. 47

(Table 3.5) Factorial ANOVA .............................................................................................. 48

(Table 3.6) The effect of different independent variables on fracture load (N). .................. 49

(Table 3.7) Incidence of different types of fracture mode for all groups. ............................ 50

(Table 3.8) Effect of different independent variables on fracture mode .............................. 51


porcelain ....................................................................................................................... 79

(Table 4.2) Firing programs for materials used in the study ................................................ 80

(Table 4.3) Two-way ANOVA ............................................................................................. 81

(Table 4.4) Fracture load (N) data for all groups. ................................................................. 82

(Table 4.5) Effect of different independent variables on fracture mode .............................. 83

(Table 4.6) Effect of different independent variables on fracture extent ............................. 84


layer porcelain ............................................................................................................ 115

(Table 5.2) Firing programs for materials used in the study .............................................. 116

(Table 5.3) Fracture load (N) and mode for all groups ....................................................... 117

(Table 5.4) Two-way ANOVA ........................................................................................... 118

(Table 6.1) Firing program for materials used in the study. ............................................... 140

ix

(Table 6.2) The chemical compositions of the different materials used in this study as

provided by the manufacturer. .................................................................................... 141

x

List of Figures

(Figure 3.1) Two of the two-piece split molds used to fabricate the specimens. ................. 52

(Figure 3.2) Porcelain- Zirconia Specimen .......................................................................... 52

(Figure 3.3) Examples of the fractured samples ................................................................... 53

(Figure 4.1) Preparation dimensions of the master die (buccal view) .................................. 85

(Figure 4.2) Crown Fabrication ............................................................................................ 86

(Figure 4.3) Examples of different fracture extent ............................................................... 87

(Figure 4.4) SEM photomontage showing a fractured crown from group 2 (fracture

involved the veneering layer only). .............................................................................. 88

(Figure 4.5) SEM picture for a crown from group 4 showing remnant of the veneering

porcelain attached to the zirconia core. ........................................................................ 89

(Figure 4.6) SEM photomontage showing a fractured crown from group 4 (fracture

involved the veneering and core layers). ...................................................................... 90

(Figure 4.7) Fracture mode observed in metal-ceramic crowns. .......................................... 91

(Figure 5.1) Preparation dimensions of the master die (buccal view) ................................ 120

(Figure 5.2) Crown fabrication ........................................................................................... 121

(Figure 5.3) Fracture modes ............................................................................................... 122

(Figure 5.4) SEM picture for EMAX crown with non-anatomical core showing some

flaws (arrow) between the core material and the fusion ceramic ............................... 123

(Figure 5.5) SEM photomontage showing a fractured EMAX crown with anatomical core

(fracture involved the veneering and core layers). ..................................................... 124

(Figure 6.1) ToF-SIMS images for sample 1 (manually-applied veneering porcelain with

liner). ........................................................................................................................... 142

(Figure 6.2) Tof SIMS image for sample 2 (manually-applied veneering porcelain

without liner) .............................................................................................................. 143

(Figure 6.3) ToF-SIMS images for sample 3 (pressed veneering porcelain with liner). .... 144

xi

(Figure 6.4) ToF-SIMS image for sample 4 (pressed veneering porcelain without liner) . 145

(Figure 6.5) Positive ToF-SIMS spectra for the IPS e.max Ceram powder (right) and

Zirliner powder (left). ................................................................................................. 146

(Figure 6.6) Changes in the relative atomic percentage as recorded by the XPS method

across the interfacial region for sample 1. .................................................................. 147







(Figure 6.10) Tof-SIMS spectra showing a comparison of the Zr and ZrO regions from

pixels close to and far away from the interface in the core area................................. 151

xii

List of Appendices

Appendix 1 ......................................................................................................................... 160

Mechanical Cyclic Fatigue Studies............................................................................................................... 161

Appendix II ......................................................................................................................... 165

Studies about the Clinical Performance of All-ceramic Crowns ..................................................... 166

Appendix III ....................................................................................................................... 179

Bond Strength of Zirconium Oxide to Veneering Porcelain ............................................................. 180

1

Chapter 1: Introduction and Literature Review

2

1.1. Introduction

The increasing demand for esthetics in combination with health and environmental

concerns about some metallic restorations stimulated dentists to explore and consider

metal-free tooth-colored restorations. Advances in the dental ceramic industry have

introduced different types of all-ceramic restorations.

All-ceramic restorations have various advantages over metallic and composite resin

restorations including:

Life-like appearance in terms of color, surface texture and translucency, 1-3

as

a result of a wide range of translucency – opacity combinations achieved with

these systems, and an acceptable well-contoured emergence profile.4

Biocompatibility: The glass-like properties render the all-ceramic restoration

biocompatible and well-tolerated by the surrounding tissues.5, 6

The surfaces of

all-ceramic restorations, if well-polished, glazed and highly smooth, reduce

dental plaque adherence.5, 7

In addition, a supra-gingival or at-the-gingiva all-

ceramic margin will not significantly jeopardize the esthetics and will reduce the

possibility of gingival involvement during tooth preparation, impression taking

and function.4

Wear resistance: Porcelains showed less wear rate compared to type 1 gold

alloys, base-metal alloys, cobalt-chromium alloys and composite resins.8

Color stability: All-ceramic restorations show better color stability in

comparison to composite resin restorations. 9

The original all-ceramic restorations have the following disadvantages:

Susceptibility to fracture.

Inadequate marginal fit and microleakage: However, developments in

ceramic formulation and fabrication techniques combined with improvements in

bonding systems and the utilization of computer technology have introduced

new all-ceramic systems with improved marginal adaptation. The use of resin

cement improved the marginal integrity and reduced the microleakage of all-

ceramic restorations. 10

3

Excessive wear of the opposing teeth: Wear has been observed in the natural

dentition opposing ceramic restorations.1 This observation received considerable

attention in the literature. However, studies in this field are affected by the lack

of universal agreement in relation to the interpretation of laboratory wear tests

and their clinical significance.11

Technique sensitivity: The clinical aspect (preparation, impression and

cementation) as well as the laboratory part for making all-ceramic restorations

need special attention.12

1.2. Dental Ceramics Classification 13

Dental ceramics can be classified according to:

1. Fusion temperature:

High-fusing (1315-1370 C)

Medium-fusing (1090-1290 C)

Low-fusing (870-1065 C)

Ultra-low-fusing (<870 C)

2. Application for:

PFM restorations.

All-ceramic restorations.

Denture teeth.

3. Fabrication Techniques:

Sintered. (e.g., Vitadur Alpha)

Slip-cast. (e.g., In-Ceram)

Heat-pressed. (e.g., IPS Empress and IPS Empress 2)

Machined. (e.g., Vita Mark II and ProCad)

4

4. Crystalline phase: Can be classified according to the nature and amount of

crystalline phase. Different crystalline phases have been used for the fabrication

of ceramics, for example:

Feldspar (KAlSi3O8) (e.g., Vita Mark II)

Leucite (KAlSi2O6) (e.g., IPS Empress)

Mica (KMg2.5Si4O10F2) (e.g., Dicor)

Alumina (Al2O3) (e.g., In-Ceram Alumina)

Spinel (MgAl2O4) (e.g., In-Ceram Spinel)

Lithium disilicate (Li2Si2O5) (e.g., IPS Empress 2)

Yttrium stabilized zirconium oxide (ZrO2) (e.g., YZ Cubes)

1.3. Indications for All-Ceramic Restorations

All-ceramic restorations are indicated in areas where maximum esthetic is required

and the occlusal forces are favorably distributed.1 They are also indicated in individuals

with allergy to dental alloys.

1.4. Fracture of All-Ceramics Materials and Restorations

The strength of a certain material can be predicted based on the strength of the bond

between its atoms. In reality, most of materials fail to exhibit the predicted strength due

to different factors. For example, the ability of the ceramic materials to support the

applied forces is compromised by the presence of defects and flaws combined with their

low fracture toughness.14

Failure begins with microscopic damage that has resulted from

interaction of preexisting defects with applied load.14

The fracture strength of ceramics is

determined by the largest and most favorably-oriented crack in relation to the applied

force. Two types of defects can be found in ceramics: fabrication defects (that arise

during processing or as microstructural features) and surface cracks (due to machining

and grinding processes).15

Due to the variation of crack sizes, the strength of ceramics

exhibits large variability and special statistical methods are required to account for this

variability (e.g Weibull distribution).

5

Clinical failure can occur due to impact forces or subcritical crack growth.16, 17

The

failure of a ceramic restoration due to impact force is very rare and might result from

trauma or sudden biting on hard object. Most ceramic clinical failures result from

subcritical crack growth, as a consequence of repetitive application of low-level loading

below the strength of the material (fatigue), which is enhanced in the aqueous

environment.18, 19

These flaws act as an area of stress concentration at their tips although the overall

stress throughout is relatively low. The stress concentration at these areas is directly

related to the applied stress and the length of the flaw and inversely related to the width

of the flaw. The reaction to the stress concentration differs between ductile and brittle

materials. In ductile materials, the stress at these flaws is reduced due the blunting of

crack tip as a consequence of the plastic deformation. On the other hand, the brittle

materials lack the ability to deform plastically, which render them more susceptible to the

stress concentration.

Beside these minute inherited flaws on the surface or the interior of this material,

other areas of high stress concentration can arise at:20

Areas of large surface or interior defects, such as porosity, inclusions,

grinding roughness and machining damage.

Areas of abrupt change in the shape of the structure (e.g. abrupt change in the

thickness of the material or sharp internal angles).

Interfaces between two materials with a large difference in the modulus of

elasticity or the thermal compatibility.

Areas where the load is applied at a point to the surface.

1.5. Fatigue Process in Ceramics Materials and Restorations

The term fatigue was first proposed by Panalet in 1839, as a response to the increase

in the use of the rapidly moving parts in industrial application. In those times, the fatigue

phenomenon was explained by the crystallization of material after continuous use, which

was thought to increase its brittleness and its susceptibility to fracture.

6

Today, fatigue failure is explained by the development of microscopic cracks in area

of stress concentration. These cracks tend to grow with time under dynamic loading

which weakens the material and might lead to catastrophic failure if the applied load

exceeds the mechanical capacity of the remaining intact portion of the material.

In metal, this phenomenon can be explained by the mobility of point defects and

dislocations. Based on this explanation, it was believed that brittle materials, such as

ceramics, are free of true cyclic fatigue effects due to the presence of strong covalent and

ionic bonds which will essentially limit the possibility of dislocation motion. This belief

was maintained for years as a result of the absence of data about the effect of cyclic

fatigue on brittle materials due to the difficulty of measuring subcritical crack growth on

these materials, in addition to the difficulty to make conclusive results about the effect of

fatigue due to the significant scatter of the stress-life data of these materials.21

Our current understanding of the fatigue process in ceramics is based on the fact that

kinematic irreversible cyclic deformation can arise not only from dislocation slip. Several

other mechanisms have been suggested including: microcracking, martensitic

transformation, creep, interfacial sliding, crazing and shear flow. The microscopic

deformation in brittle material can arise due to one of the following mechanisms: 21

1. Frictional sliding along the faces of microcracks.

2. Wedging of microcracks by debris particles.

3. Inelastic strain generated by dilatation and shear transformation.

4. Viscous flow of glass phases.

5. Creep cavitations.

6. Fracture of the reinforcement phase.

The mechanism of the cyclic crack growth in ceramic can be classified into 2

classes:21

Intrinsic mechanism: in which the crack advancement results from

microscopic deformation at the crack-tip in response to cyclic loading.

7

Extrinsic mechanism: in which the crack advancement resembles the

advancement in the monotonic loading, and the unloading process enhances the

growth due to the decrease in the effect of crack-tip shielding.

In ceramics, a crack can become unstable under static stress alone, in the absence of

cyclic loading, in a phenomenon termed static fatigue, which is related to the chemical

interaction of water with silica at the crack tip. The presence of the water molecule will

cause the breakage of ester bond on the silica and the development of two hydroxide

molecules. This effect is obvious in ceramics with high-silica content. Water is available

for the external surface of the restoration from the saliva while both saliva and dentinal

fluid can reach the internal surface of the restoration through the cement.4

1.5.1. Factors Affecting the Fracture and Fatigue of Ceramic Restorations

In ceramic specimens that are not bonded to a supporting structure, the fracture

originates from flaws or defects on the tensile side directly opposite to the compressively

loaded area when specimens loaded in flexure. In the clinical situation, the monolithic

crown is supported by a material with lower modulus of elasticity and is often bonded to

it.22

This system builds more complex stresses at the ceramic-cement interface directly

below the loading area. These stresses arise from the mismatch in the produced

deformation due to mismatch in the modulus of elasticity between ceramic, cement and

supporting structure. These stresses can be seen in the finite element analyses.23, 24

In

multilayer ceramic systems, different ceramic materials add more complexity to the

interfacial stresses. These interfacial stresses are more sensitive to the ratio of modulus of

elasticity but less sensitive to the thickness ratios of ceramic, cement and supporting

structure.22

Several factors might affect the fracture strength of all ceramic crowns in-vivo:

Factors related to the restoration:

1. Composition of ceramic material: The mechanical properties of the

material are affected by the nature, amount, size and distribution of the

crystalline phase as well as the mechanical properties of the glassy matrix.25,

26

8

2. Processing: Microstructural flaws and defects (e.g. porosity, microcracks

and impurities) can develop in the ceramic restoration as a result of

inaccurate or suboptimal processing. 27

3. Finishing and glazing affect the ceramic texture and roughness.28-32

4. Crown dimensions and geometry: the fracture resistance of an all-

ceramic crown will increase if the crown thickness is increased.33, 34

5. Core/veneer thickness ratio (in multilayer restoration): Increasing the

core/veneer thickness ratio increases the flexure strength and shifts the

crack initiation site from veneer to core.35

White et al.36

reported that

increasing the core/veneer thickness ratio increases the modulus of rupture

of beams that were composed of zirconia core and matchable veneering

porcelain materials. Fleming et al. 37

found similar results, an increase in the

number of fracture fragments, Hertzian cone and delamination were

associated with a decrease in core/veneer thickness ratio.

6. Differences in ceramic mechanical and physical properties (modulus of

elasticity and thermal contraction coefficients) in a multilayer restoration.4

o Differences in the coefficient of thermal expansion (CTE) between

the materials: Mismatch in the CTE can induce residual stresses that

can cause immediate or delayed failure. For the metal ceramic crowns,

the veneering porcelain should always be under slight compression.

These compressive stresses increase the strength of the restoration by

inhibiting the crack propagation by keeping them closed.3 This can be

accomplished by choosing a veneering porcelain material with a CTE

that is slightly lower than that of the metal.38

In an all-ceramic

restoration, this mismatch in the CTE will introduce tensile stresses in

the ceramic core which might have a negative effect on some weak

brittle core materials because the tensile strength of a brittle material is

much lower than its compressive strength.39

Steiner et al. reported that

a mismatch in the CTE value between a ceramic core and veneering

porcelain less than 1 ppm/C does not produce visible cracks in the

multilayer all-ceramic restoration.40

Isgro et al. stated that the linear

9

CTE measured according to the ISO standard for metal-ceramic

systems was not appropriate as a predictor for the thermal

incompatibility for all-ceramic systems because the ceramic materials

showed non-linear expansion.41

Equal CTE for the core and the

veneering porcelain is not enough to predict thermal compatibility.

Other factors should be considered such as the visco-elastic behavior of

porcelain, the effect of multiple firing and the effect of cooling rate.42

Studies have confirmed that leucite-reinforced veneering porcelain

showed a change in the CTE with each heat treatment. This could be

explained by the change in the leucite content after multiple firing43

and

the de-coupling of the leucite crystals from the glass matrix during

cooling and re-coupling to the glass matrix with firing.44

o Differences in modulus of elasticity of the different layers:

Increasing the modulus of elasticity ratio (E veneer / E core) will increase

the critical load needed to cause radial cracks in the core. This will be

beneficial up to a certain limit after which using a too stiff veneering

material will increase the susceptibility of radial cracking in the

veneering layer.45

7. Internal fit of the restoration: To prevent crown fracture, the crown must

be fabricated with the best possible fit.46, 47

Factors related to the supporting structure (natural teeth or other core

materials) of the restoration:

1. Modulus of elasticity: The fracture strength of all-ceramic crowns will

increase if the elastic modulus of the supporting structure is increased.48

2. Preparation characteristics (preparation geometry, dimensions and

geometry):

o Increasing the preparation dimensions will increase the fracture

strength of the crowns. This effect is obvious in comparing the fracture

strength of molars with premolars and anterior crowns.

o Increasing the occlusal thickness of the crown as a result of

decreasing the preparation height will increase the fracture strength of

10

the crown.49

On the other hand, this approach is not conservative and

tends to affect the retention form of the preparation.

o A total occlusal convergence of 10º provides the best combination

between fracture strength and conservation of tooth structure. A

convergence angle of five degrees is associated with the lowest fracture

strength, and a 15 º or 20º convergence angle need greater tooth

reduction.49

o Doyle et al. found an increase in the fracture strength of crowns

prepared with shoulder finish line and sharp axiogingival line angle

compared to a shoulder with round axiogingival line angle and a

chamfer.49

In contrast, Burke et al. reported that preparing a 1mm

shoulder finish line rather than a knife edged margin did not have a

significant effect on the strength of all-ceramic crowns.50

Factors related to the oral environment surrounding the restoration:

1. Oral fluids: An aqueous medium will enhance crack growth and decrease

ceramic restoration strength in a phenomenon called “static fatigue”.19

2. Changes in temperature: Restorations in the oral cavity are subjected to

thermal changes which may cause stress concentration at the restoration

cement interface due to mismatch in the CTE of the ceramic, cement and

supporting structure.

3. Occlusal forces: Clinically, dental restorations are subjected to cyclic

forces. These cyclic forces have both vertical and lateral components,51

range between 60 and 250 N during function and could reach 500 to 800 N

for short periods of time.14

The range of forces in maximum biting is 400-

890 N in the molar region, 222-445 N in the premolar area, 133-334 N on

the cuspids and 89-111 N on incisors.52

In the posterior teeth, these forces

are distributed on an average contact area of 52 and 64 mm2

for females

and males, respectively.51

In an average individual, the number of cycles of

mastication per day ranges between 800 -1400 cycles14

and can reach up to

2700 cycles/day.53

This translates to roughly 290,000 - 106

cycles/year.

This number should be reduced by factor ranging from 5 to 20 because not

11

all the chewing cycles are active (i.e. representing maximum load in the

chewing cycle).53

Factors related to the cementation of the restorations:

1. Adhesion mechanism: An increase in the fracture strength has been

reported with the use of adhesive cements compared with non-adhesive

ones.54, 55

2. Cement thickness: Increasing the cement thickness reduces the fracture

strength of all-ceramic crowns.47

3. Mechanical properties of the cement (mainly modulus of elasticity): As

the elastic modulus of the cement increases, the crown-cement-supporting

structure system becomes stiffer. This decreases stresses in the crown.56

An

exception to this, however, occurs when resin cements (low modulus) are

compared to zinc phosphate cements (high modulus) for the cementation of

all-ceramic crowns.54,55

This conflict can be explained by the beneficial

effect of the adhesive cements on the fracture strength of all-ceramic

crowns.

4. Efficacy of the bond between the cement and porcelain: This is affected

by the treatment of the internal surface of the crown.

5. Efficacy of the bond between the cement and supporting structure.

Rekow et al. 56

used finite element analysis to conduct a factorial analysis of some

variables influencing stresses in all-ceramic crowns. They studied the effect of seven

factors (crown material, crown thickness, cuspal inclination, cement elastic modulus,

cement thickness, supporting tooth core, and location of occlusal loading). Crown

material and thickness accounted for 58.2% and 24.7% of the variability in the principal

stresses in the crown, respectively. Elastic modulus of the cement, supporting tooth core

and loading position accounted for 2.1%, 0.5% and 1 % of the variability in the stresses

respectively. Sensitivity to these factors may not be the same for crowns made with

different materials.

12

1.5.2. Improving the Strength of Ceramic Material

Different methods have been proposed to improve the strength of the ceramic

restoration:57

1. Method of strengthening the material:

Removal of surface flaws or reducing their sizes and numbers by

improving the processing methods, polishing and glazing.

Development of residual compressive stresses: The development of

residual compressive stress on the surface of the material will help in

improving the strength of the material. Strengthening will arise from the

fact that these compressive stresses should be overcome first before the

development of any net tensile stresses. Different methods could be used to

develop residual compressive stresses:

o Ion exchange: The aim of this method is to change the ion in the

surface by larger ions, which should be squeezed on the place

previously occupied by the small ion. This squeezing will lead to the

development of residual compressive stresses on the surface.

o Thermal tempering: This technique creates residual compressive

stresses by rapidly cooling the surface of the material while the material

is hot and softened. This will result in development of residual

compressive stresses on the surface and residual tensile stresses on the

center.

o Thermal compatibility: The use of veneering porcelain on the

surface of the crown with a slightly lower coefficient of thermal

expansion (CTE) compared to the core material will result in

development of residual compressive stresses at the surface of the

veneering porcelain. However, such a mismatch in the CTE might

result in the development of residual tensile stresses at the interface

between the two materials, which might weaken the bond between

them.

13

Interruption of crack propagation: This technique requires the

incorporation of different material inside the ceramic capable of hindering

the cracks. Different methods can be used:

o Dispersion of a tough crystalline material to disrupt the crack

propagation.

o Transforming toughness: This method is obvious for the yttrium-

stabilized zirconium oxide ceramics (Y-TZP). Tetragonal crystals in Y-

TZP are metastable and can be transformed to larger monoclinic

crystals with the application of any stress from cracks or flaws 58

. This

phenomenon is beneficial in hindering crack growth and increasing the

fracture toughness and is therefore called “transformation

toughening”.59

2. Designing structures to minimize the stress concentration and tensile stresses.

1.5.3. Laboratory Testing of the Performance of the Ceramic Materials and Restorations

For years, the focus of most of research involving the ceramic materials was directed

toward evaluating the strength of these materials under static loading by measuring the

strength of either ceramic standard specimens or ceramic restorations.

Many studies reported on strength properties of different ceramic materials utilizing

different tests (three-point bending, four-point bending and biaxial flexure test). It is

obvious that the test type affects the strength data. Even with the same test, the strength is

affected by the test methodology and specimen geometrical characteristics. These

strength data can be used for comparison between materials within the same testing

methodology. However, strength data should not be considered as “absolute” values and

extrapolating in-vitro strength data to the clinical performance must be considered

cautiously.11

A number of studies reported the in-vitro fracture strengths of different ceramic

crowns and fixed partial dentures (FPDs). Data from these studies were not consistent for

the same material because the results were affected by the testing methodology. Kelly22

criticized the load-to-failure tests of crowns and FPD’s and questioned their clinical

relevancy. He stated that data collected from traditional load-to-failure tests are not valid

for the following reasons:

14

1. The damage reported in in-vitro studies start from the occlusal surface

whereas clinical failures of monolithic restorations start from flaws and stresses

at the cementation surface.

2. Extremely high fracture loads (>1500 N) are required in-vitro compared to

the smaller occlusal loads measured in the oral cavity during function.

3. Incorrect stress state: The load applicators, used in most of the in-vitro

studies, tend to produce higher and incorrect stress distribution compared to the

clinical situation.

4. It is common to see multiple piece fractures in traditional load-to-failure tests

compared to single fracture in clinically failed crowns.

5. For in-vitro studies, fracture loads are recorded with a sharp drop in the load.

However, the “pop-in” crack on the cementation surface is not accompanied by

any drop in the load because the loading piston is entirely supported by the

intact crown. One method to record the “pop-in” crack reported in the literature

is the use of acoustic monitoring.60

During the last 2 decades, the trend of the dental material research shifted toward

subjecting the ceramic standard specimens or restoration to cyclic loading to simulate the

clinical situation. Both the effects of thermal and/or mechanical cycling on strength have

been studied for different ceramic materials. Many of these studies reported a decrease in

the strength after thermal and/or mechanical cycling.29, 61-67

1.5.4. Methods for Testing the Effect of Fatigue on Ceramic Materials and Restorations

Several approaches have been used to evaluate the effect of fatigue on the dental

ceramics:

1. Phenomenological lifetime data (fail or not-fail approach)

In this approach, a standard specimen or surrogate structure (e.g. crown)

will be subjected to cyclic fatiguing using different loads until fracture. The

number of cycles (N) to fracture is recorded and plotted against the stresses (S)

to get the S-N curve. For simplicity the S-N curve can be divided into 3

regions:53

15

Low cycle fatigue (ranges between 1 and 10 4

cycles): in this region, the

applied stresses are often higher than the elastic limit (i.e will cause plastic

deformation in the material). Therefore, this type of fatigue has limited

application in testing the ceramic restoration.

Limited endurance fatigue (ranges between 10 4

and 10 7

cycles): the

stresses utilized in this region lie within the elastic regimes of the material.

Unlimited endurance fatigue (above 10 7

cycles).

Theoretically, it is logical to design the ceramic restorations with optimum

mechanical-fatigue resistance. However, clinically restorations may fail due to reason

other than mechanically-based ones (e.g. biological failure), which might limit their

survival. For this reason, in conjunction with the financial difficulty in conducting cyclic

fatigue tests for high numbers of cycles (especially, if combined with low frequency to

resemble the intra-oral loading), it is suggested to conduct the conventional endurance

limit of the material or restoration. This conventional endurance limit can be defined as

the stress amplitude that 50% of the sample will sustain for a predetermined number of

cycles (between 10 6

and 10 8

cycles).53

Similar to strength, fatigue strength is affected by

the distribution of the flaws in the sample. Accordingly, it is normal to have dispersion of

the stress values at a predetermined number of cycles.

It is not uncommon for some of the materials to survive the cyclic loading for a

predetermined number of cycles using a specific predetermined load. Alternatively, the

specimen might be loaded to failure after the cyclic loading and the fracture load or

fracture stress will be recorded. Obviously, this approach will not give enough data about

the fatigue life of the sample other than the fact that the specimen or the structure

survived the fatigue condition without failure, but it might help us in comparing the

effect of the cyclic loading on the strength of the different tested materials. In fact, most

of the studies, which evaluate the effect of fatigue on dental ceramic materials and

restorations, utilized this alternative approach.

Another method of fatigue testing is the “staircase” approach in which the crown or

specimen is subjected to a specific load for a predetermined number of cycles. If the

crown fails, the load is decreased by a certain value (one step) and another sample is

loaded for the same predetermined number of cycles. If it does not fail, the load is raised

16

by the step size and another crown is loaded for the predetermined number of cycles and

so on. This approach will allow us to calculate the fatigue limit and its SD (Standard

Deviation).68

At least 20 samples should be loaded to measure the fatigue limit

appropriately.53

This approach is not conservative and is not useful for the prediction of

lifetime. Also, the applicability of the results is sensitive to the selected number of cycles,

since the fatigue for a low number of cycles will preclude observations at the high

number of cycles. Another disadvantage of this approach is the cross-over behavior, in

which some material will perform better at low stresses level, while others might perform

better at high stresses level.68

2. Crack-growth data:

Two different approaches have been utilized to study the crack growth

behavior in ceramics:

Monitoring the rate of crack advancement in response to cyclic fatigue

using standard ceramic specimens (with or without pre-crack). Eventually, a

plot of the crack growth rate per cycle (da/dN) vs. the difference in the

stress intensity factor (∆K) will be generated. This approach needs a

specific specimen design and a frequent measuring of the crack length at

periodic intervals.

Difference in the stress intensity factor can be calculated using the

following equation ∆K= Kmax-Kmin = Y ∆σ √πa

Where, ∆σ (difference in the stress) = σ max- σ min , a is the crack length and

Y is geometrical factor.

The general shape of this plot has been established by Paris and his

coworkers. Three different regions can be distinguished in this graph:

o Region I: in this region, the fissure propagates in crystallographic

shear mode. The crack growth in this region is mainly mediated by

water, especially in ceramics with appreciable silica content.

o Region II: The effect of moisture is less pronounced in this region

and the direction of the crack growth will be normal to the tensile stress

(plane strain condition) and will be 45 degrees to the stress axis if the

17

stress at the crack tip is significantly influenced by the specimen

geometry.

o Region III: Catastrophic failure occurs in that region by

intergranular cleavage

Measuring the crack growth parameter using the dynamic fatigue test. In

this approach, the strength of the material is determined as function of the

stressing rate. The specimens used in the test may contain either an

indentation-induced flaw of certain size or the inherited flaws only.

The collected strength data will be used to estimate the fatigue parameter

using the following equation: 69

s f = B n +1( )s i

n-2s ·[ ]1/n+1

where

s iis the inert strength (inherent, non-fatigue stress at failure,

usually measured in oil to preclude the effect of the water),

s · is the

stressing rate and B is a parameter associated with A, n, fracture toughness,

crack geometry and loading configuration. The slow crack growth

parameter n and fatigue parameter B can be determined from a plot of ln

s f

as a function of ln

s · by linear regression of the data with the following

equation:

lns f =1

n +1lns · + lnb

where

b =1

n +1lnB+ ln n +1( ) + (n - 2)lns i[ ]

This method is preferred compared to the fracture-mechanics based

crack growth tests, because it gives more conservative life time predication

and because the flaws (in case of using specimens with inherent flaws only)

are more realistic and simulate the clinical situation.

There are some disadvantages in utilizing this technique:

o Difficulty in getting the inert strength of the material.

18

o Difficulty in choosing the range of the stressing rates due to

inconsistency in the response of different materials to stressing rate (i.e

some materials show a drop in the strength at high stressing rate).

o The fatigue parameters measured using this method do not

correspond with the parameters collected using the repetitive cyclic

loading.

3. Contact fatigue data: This technique is used to study the damage mode after

cyclic loading using a sharp or blunt indenter. The applicability of this technique

for studying the ceramic material has been limited previously due to the

difficulty encountered in monitoring and measuring the different damage modes

on relatively opaque dental ceramics, which required sectioning of different

samples at different stages during the test and examining them. Recently, a new

route was used, in which a transparent material with mechanical properties

matching those of the ceramic, cement and tooth structure were used to observe

the crack initiation and propagation without sectioning the samples.

1.5.5. Simulating the Clinical Situation in the Fatigue Testing

Some of the available mechanical cyclic fatigue tests have been reviewed and

revealed no consistent protocol. The test parameters (material, size and shape of the load

applicator, frequency, range of force, and number of cycles) have been varied in these

tests (Appendix I).

To simulate the clinical situation the following should be controlled:

1. Factors related to the loading environment: The test should be performed in

the water and preferably under thermocycling to resemble the oral environment.

22

2. Factors related to the sample:22

Samples should be fabricated to have realistic flaws that resemble the

clinical situation.

Reliable adhesive cements should be selected, with a narrow range of

elastic modulus.

19

It is difficult to standardize the dimensions and modulus properties of

natural teeth. Therefore, the use of an alternative with a modulus of

elasticity resembling dentin will help standardize the different variables

related to the crown’s supporting structure.

3. Factors related to the applied load:

Magnitude: A controversy can be seen in the literature regarding whether

the load should be selected according to the average biting force for an

average individual or according to the worst-case scenario by applying a

load comparable to the maximum biting force that was measured. Another

factor that could govern the force selection is the location of the tested

restoration because as stated previously, the biting forces increase from the

anterior to the posterior region.

The contact pressure applied should fall within the range of the measured

contact pressure clinically. The contact pressure depends on the contact area

between the load applicator and the sample and to lesser extent, on the

elastic modulus of the load applicator. 22

Force direction: It is preferable to subject the samples to both vertical

and lateral forces. The applicability of this suggestion is determined by the

loading machine capability to perform the cyclic loading in both directions.

Number of cycles: 10 6

and 10 8

cycles seems to be reasonable for testing

the dental ceramic materials. 53

Frequency of the cyclic loading: It is advisable to use a frequency

comparable to the frequency of the chewing cycle (1-2 Hz). This frequency,

if combined with a high number of cycles as suggested above, requires a

long period to perform the test and increases the cost of the test. 53

1.6. Survival Rates of All-Ceramic Crowns

Twenty-two clinical studies were included in a systematic review about clinical

complications in fixed prosthodontics.70

All-ceramic crowns showed the lowest

prevalence of complications (8%). However, most of these studies were of short-term

20

duration. In 18 of them, the study duration ranged between 1-4 years while 4 were for

more than 5 years. The most common complications encountered were crown fracture

(7%), loss of retention (2%), and need for endodontic treatment (1%). The incidence of

crown fracture increased with the length of the study. The relation between the fracture

incidence and location in the arch was discussed in ten articles. The molars showed

higher fracture rate (21%) compared to premolars (7%) and anterior teeth (3%).

Kelly 4 reviewed ten articles about the survival rates of all-ceramic crowns. He found

that four materials (IPS Empress, In-Ceram Alumina and Spinell, and Procera Allceram)

have been extensively studied. He stated that the lowest failure rates were for In-Ceram

Alumina and Procera Allceram. The fracture rates seemed to be higher in the posterior

crowns compared to the anterior.

Studies that have reported the survival and fracture rates of all-ceramic crowns are

summarized in Appendix II.

1.7. Zirconium Oxide Ceramics

Zirconium oxide occurs in three crystalline forms: monoclinic, tetragonal and cubic.

Pure zirconium oxide occurs as monoclinic crystals at room temperature and transforms

into tetragonal at 1170ºC and then to cubic crystal at 2370ºC. During cooling, the crystals

will transform to monoclinic crystals at about 970 ºC. This transformation is

accompanied by a decrease in the flexure strength and volume expansion (3-4%).71

Zirconium oxide cubic or tetragonal crystals can be partially stabilized at room

temperature by adding a stabilizing oxide (e.g., CaO, MgO, CeO2 or Y203). Addition of

yttrium oxide to pure zirconia forms either PSZ (Partially Stabilized Zirconia) or Y-TZP

(Yttrium Tetragonal Zirconia Polycrystals). At room temperature, PSZ consists mainly of

cubic crystals with traces of monoclinic and tetragonal crystals, while Y-TZP consist of

tetragonal crystals. The amount of tetragonal crystals depends on grain size, yttrium

oxide content and the degree of constraints exerted on them by the matrix.71

Tetragonal

crystals in Y-TZP are metastable and can be transformed to larger monoclinic crystals

with the application of any stress from cracks or flaws.58

This phenomenon is beneficial

in hindering crack growth and increasing the fracture toughness and is therefore called

“transformation toughening”.59

21

The aging of Y-TZP at 230 ºC for 400 hours in a water vapor environment caused a

reduction in the flexural strength.72

This reduction in strength was reversed with 1,000 ºC

reheating for 24 hours and attributed to micro-crack formation at the surface as a result of

tetragonal-monoclinic crystal transformation. Ishgi et al. 73

reported an increase in the

flexural strength of polished or sandblasted zirconium oxide ceramic after sintering

compared to control sintered zirconium oxide specimens. Also, they reported a reduction

in the flexural strength of the ground samples after “veneer simulated firing”. This

reduction in the flexural strength after heat treatment could be explained by relaxation of

the residual compressive stresses that developed on the surface of the specimens due to

surface treatment.74

Papanagiotou et al.75

found no significant degradation of the flexural

strength of zirconium oxide samples subjected to polishing and low-temperature

degradation (24 hours or 7 days in boiling water or in humidifier at 250 ºC), while

samples treated with airborne-particle abrasion showed an increase in the flexural

strength. However, low-temperature degradation resulted in loss of yttrium, which could

decrease tetragonal-phase stability and long-term performance of zirconium.

In dentistry, Y-TZP is used to fabricate a substructure for crowns and FPDs. Y-TZP

sintering is accompanied by high volume shrinkage (20 – 25 %).58, 76

Therefore, the

current technology does not allow us to fabricate crowns and FPDs frameworks by direct

sintering on dies. There are two approaches to compensate for this sintering shrinkage in

order to develop prosthetic frameworks using partially-sintered zirconium oxide ceramics.

The first approach is to fabricate an oversized framework which will shrink to the proper

size after sintering. In this approach partially-sintered bar coded blocks are used to mill

the substructure. These bar codes are indicators of each block’s density and are used to

calculate the size needed to compensate for the actual sintering shrinkage. The other

approach is to fabricate the framework on an oversized die. After sintering, this oversized

framework will shrink to the proper size. Both approaches need precise measurement of

the sintering shrinkage utilizing the CAD/CAM technology. Another approach is to grind

the crowns using fully sintered zirconium oxide.77

Because it is fully sintered, the

infrastructures are milled to the exact size. The milling process of fully sintered Y-TZP is

slower, increases the wear of the cutting hardware and can introduce some microcracks

in the core material.58, 78

22

Several Y-TZP systems are available for the fabrication of the infrastructure for

crowns and FPDs. The majority of these infrastructures are fabricated from partially

sintered Y-TZP blanks (e.g., In-Ceram YZ, Vita Zahnfabrik, Bad Sackingen, Germany;

Cercon, Dentsply Ceramco, York, PA; Lava, 3M Espe, St. Paul, MN, Procera AllZirkon,

Noble Biocare, Yorba Linda, CA). Fully sintered Y-TZP can be used to fabricate the

infrastructures in some of the systems (e.g., DC-Zirkon, Smartfit, Austenal, Chicago,

IL).76

One of the disadvantages of Y-TZP is that it is not transparent and can not be stained

to give good esthetic results. Therefore, Y-TZP must be veneered with suitable veneering

porcelain to enhance esthetics.36

The application of the veneering porcelain on top of the ceramic coping could

introduce multiple flaws and interfaces, and complicate the nature of stresses within the

restoration. In a previous study, we reported that the fracture of the crowns made with

zirconium oxide substructure was a combination of chipping within the veneering

porcelain and delamination at the core-veneer interface.79

Interfacial failure was the

basic cause of failure for Y-TZP crowns in other in-vitro studies.80-82

Data collected from

failed clinical crowns showed a similar type of failure for multilayer crowns.83

In general, the fracture of multilayer crowns starts at their weakest part. In cases

where a stronger and stiffer core substructure is veneered with weaker porcelain, the

failure typically occurs in the weak veneering porcelain or at the weak bond between the

core and veneer.84

Several factors can affect the flexural strength of all-ceramic

veneering materials. The effect of these factors has been discussed previously. A weak

core-veneer bond can result from:

1. Structural flaws at the interface that might develop during veneering

porcelain application due to poor wettability.85, 86

2. Tetragonal monoclinic transformation which causes microscopic uplifts on

the surface due to the formation of larger monoclinic crystals.87, 88

This

transformation can be initiated during the multiple firing of veneering porcelain

on the zirconium oxide core.74

3. Firing shrinkage of the veneering porcelain.84

23

4. Interfacial tensile stresses at the core-veneer interface as a result of mismatch

in CTE.85

5. Low interfacial toughness compared to fracture toughness of core material:

The ratio of interfacial toughness to the core’s toughness, as well as the elastic

modulus ratio of the two materials, are determinant factors for the propagation

of cracks along or through the interface.89, 90

Several studies reported the core-veneer bond strengths for different zirconium oxide

systems (Appendix III). The following points were noticed from these studies:

Different methodologies were used in these studies, which hinder the

comparison between the results.

Zirconia/veneer bond strength was compared to metal/veneer bond strength

in 6 studies.91-96

Two studies reported no significant difference between bond of

veneering porcelain to zirconia or metal91, 92

and 3 studies reported higher bond

strength to metal in comparison to zirconia.93-95

On the other hand, only one

study reported a higher bond strength to zirconia in comparison to metal.96

All of the studied systems showed cohesive fractures of the veneers,

interfacial core-veneer fractures or a combination of both. None of the samples

failed cohesively through the zirconia core.

The effect of the veneering porcelain composition on the bond strength did

not receive adequate attention in the literature. One study reported that the

presence of zirconium in the veneering porcelain composition improved the

bond strength to zirconia.97

Another study reported higher zirconia/porcelain

bond strength for leucite containing pressable veneering porcelains compared to

non-leucite containing pressable veneering porcelains. 98

The use of liners affected the bond strength and the fracture mode of the

tested specimens. The effect of liner on the bond strength was material

dependent. 84, 99, 100

The effect of different surface treatments was reported in some of these

studies. The results of these studies were heterogeneous and were material-

dependent for some of them.97, 99-104

The effect of different surface treatment on

24

the crystal transformation was investigated. The highest amount of the

monoclinic phase was found after sandblasting. However, these phase changes

did not influence the core-veneer bonding. 101

Five studies evaluated the bond strength of pressed veneering porcelain to

zirconia. 95,

105-107,

108

Three of these studies reported no significant difference

between the layered and pressed sample,105-107

while the results of two other

studies reported higher bond strength for the pressed samples in comparison to

the layered ones.95, 108

In addition, the use of pressable porcelain on zirconium

oxide cores shifted the failure mode to be primarily cohesive failure in the

veneering layer compared to layered veneering porcelains which showed high

percentage of interfacial failures. This improvement in the bonding strength

could be explained by the improvement in the contact between the core and

veneering porcelain materials due to applied pressure during pressing, by

decreasing the chance of microgap formation as result of deformation and

cooling stresses, and by the reduction of the residual stresses and the tetragonal

to monoclinic crystal transformation at the interface due to the elimination of the

multiple firing of the layered veneering porcelain.84

Currently, little is known about the mechanism of the bonding between

zirconia cores and the veneering porcelain. A recent study reported that zirconia

and alumina as core materials demonstrated adhesive bonding (mechanical) with

silicate porcelain veneers, whereas glass infiltrated alumina and lithium

disilicate glass ceramic have reactive bonding (chemical) towards silicate

veneers. 109

The chemical structure at the core-veneer interface was analyzed

using Energy Dispersive X-ray (EDX), which revealed that some of the veneer

elements diffused into the zirconium oxide layer to a depth of 8 -10 μm.84

It is well known that surface contamination of the zirconia core can affect its

bonding strength to the veneering porcelain. A recent study reported on

contamination of the surface of the zirconia crowns in dental laboratory. High

values of aluminum found on surface were probably caused by diffusion of

aluminum oxide from the base during sintering. This difference between

declared composition and composition found in this investigation could affect

the bonding of zirconia and layered ceramic. 110

25

During the last 3 years, many manufacturers developed colored zirconia

cores. The effect of using these colored cores on the bond strength has been

investigated. The bond strength to colored zirconia was significantly weaker

compared to white zirconia frameworks. Although no marked chemical

differences between the examined zirconia materials could be found, there were

microstructural differences, especially between white and colored zirconia and

for different zirconia frameworks from different manufacturers, which

significantly affected core–veneer bond strength values.99

The effect of mismatch in the coefficients of thermal expansion (CTE)

between the veneering porcelain and core material on the core-veneer bond

strength was investigated. Significant discrepancies in CTE between veneering

porcelains and zirconia significantly affect their bond strength and caused

spontaneous debonding during the porcelain firing. Fischer et al. 111

reported

that the mismatch in the CTE and the glass transition temperature significantly

affected the bond strength of veneering porcelain to zirconia and suggested that

a veneering ceramic for Y-TZP should have a value of ΔαΔT≈1000x10−6

(where

Δα=the difference between the CTE of the zirconia and the veneering porcelain

and ΔT=the difference between the glass transition temperature and the room

temperature) in order to provide highest bond strength of zirconia/veneering

ceramic composites.

The effect of cooling rate after veneering porcelain sintering on the

zirconia/veneer bond strength was evaluated in 2 studies. 112, 113

The first study

reported that the bond strength increased with rapid cooling in comparison to

slow cooling.113

The other study reported a material-dependent effect. The

zirconia/veneer bond strength of veneering porcelain that contains leucite

crystals in their composition was not affected by cooling rate, while samples

veneered with amorphous veneering porcelain fired according to a slow cooling

protocol showed a reduced bonding strength in comparison to samples fired

according to the fast cooling protocol. 112

The effect of multiple firing of the veneering porcelain on the bond strength

was examined by two studies. 103, 114

The results of both studies were

contradictory with one study reported an increase of the bond strength with the

26

increase in the number of firing cycles, 103

and the other study reporting a

decrease in the bond strength with the higher number of firing cycles.114

Different SEM examinations of failed zirconium oxide crowns in a previous in-vitro

study revealed multiple voids within the manually applied veneering porcelain, which

develop as a result of the human factor in building up these crowns79

. As stated

previously, these voids could compromise the ability of the all-ceramic restoration to

support the applied forces. Elimination of the human factor in application of the

veneering porcelain by the use of pressable ceramic on top of zirconium oxide

substructure is expected to improve the homogeneity within the veneering porcelain and

hence enhance its ability to withstand the applied forces.

A thorough literature search revealed nine clinical studies which reported the

survival and fracture rates of all-ceramic crowns made using zirconium oxide cores115-123

and one randomized clinical trial which compared the survival and fracture rate of full

coverage zirconium-oxide crowns to full coverage gold crowns124

(Appendix II). Only

two of these nine studies were randomized clinical trials that compared all-ceramic

zirconium-oxide crowns to crowns made using glass-infiltrated zirconia-reinforced

alumina core115

or metal-ceramic crowns118

. The rest of the studies were case series. The

number of zirconia-based restorations ranged between 15 and 248 crowns and eight

different zirconium-oxide systems were used in these studies. The follow-up periods

ranged between 1 and 5 years. All the studies, except two, reported at least an incidence

of veneering porcelain fracture during the follow-up period. Most of these studies did not

report any details about the coping design and the firing protocol of the veneering

porcelain.

27

Chapter 2: Rationale and Objectives

28

Rationale and Objectives

The previous literature review makes it clear that the veneering porcelain or the

core-veneer interface represents the weakest part of all-ceramic crowns made using

zirconium oxide cores. The aim of this project was to study different factors affecting the

fatigue resistance and fracture strength of all-ceramic molar crowns made using

zirconium oxide cores.

The objectives of this thesis were:

To evaluate the effect of varying the total thickness, core thickness, and

veneering porcelain application techniques (pressing vs. manual build-up) on the

fracture strength and mode of bilayered zirconia-porcelain samples. (Chapter 3)

To evaluate the influence of varying the core thickness and modifying the

core design on the fracture strength and mode of zirconia-based crowns and to

compare these results to those of metal-ceramic crowns. (Chapter 4)

To evaluate the influence of using different veneering techniques and core

designs on the fracture strength and mode of zirconia-based crowns made using

CAD-CAM-produced veneering layers. (Chapter 5)

To study the nature of the bond between the zirconium oxide core and

matching veneering porcelain materials. (Chapter 6)

29

Chapter 3: Manuscript 1

30

Effect of total thickness, core thickness and veneering porcelain application

technique on fracture of zirconia/porcelain combinations

Mohammed H. Zahran, BDS, MSc, PhD-candidate

Department of Prosthodontics, Faculty of Dentistry, University of Toronto, Toronto, ON, Canada

Department of Fixed Prosthodontics, Faculty of Dentistry, King Abdulaziz University, Jeddah,

Saudi Arabia

*Omar El-Mowafy, BDS, PhD, FADM

Professor and Head, Department of Restorative Dentistry, Faculty of Dentistry, University of

Toronto, Toronto, ON, Canada

Asbjorn Jokstad, DDS, PhD

Professor and Head, Department of Prosthodontics, Faculty of Dentistry, University of Toronto,

Toronto, ON, Canada

University of Tromso, Norway

Amin S. Rizkalla, P Eng, PhD

Associate Professor,, Dentistry, Schulich School of Medicine & Dentistry, Western University,

London, ON, Canada

Laura E. Tam, DDS, MSc

Professor, Department of Restorative Dentistry, Faculty of Dentistry, University of Toronto,

Toronto, ON, Canada

*Corresponding author

Omar El-Mowafy, BDS, PhD, FADM

Department of Clinical Sciences

Faculty of Dentistry, University of Toronto

124 Edward Street, Toronto, Ontario M5G 1G6, Canada.

Phone: (416) 979-4934 X 4572

Fax: (416) 979-4936

E-mail: [email protected]

Short Title: Fracture of zirconia/porcelain combinations

This work was previously presented in part at the 89th General Session & Exhibition of the

IADR, San Diego, CA, USA in 2011

31

Abstract

Objective: to determine the effect of varying total thickness, core thickness and veneering

porcelain application technique on the fracture strength of bilayered porcelain-zirconia discs.

Methods: Zirconia square specimens (8 x 8 mm) were prepared in three thicknesses (0.6, 0.8 and

1.0 mm). The zirconia specimens were veneered with pressed (P) or manually-applied (M)

veneering porcelain to produce specimens with 1.5 or 2 mm total thickness. The discs were

cemented to epoxy resin blocks and loaded to failure at the center of the veneering porcelain.

Fracture load and mode were recorded and analyzed using factorial ANOVA and Chi-square test,

respectively. Results: Factorial ANOVA revealed a statistically-significant main effect on mean

fracture loads (P<0.05) for the following variables: total thickness [Mean(SD): 1.5

mm=1,178(273) N and 2 mm=1,464 (331) N], core thickness [(Mean(SD): 0.6 mm=1,196(381)

N, 0.8 mm=1,332(296) N and 1mm=1,436 (280) N] and veneering material application method

[(Mean(SD): P=1,405(310) N and M=1,237 (339) N]. Higher fracture loads were associated with

the greater total thickness, the greater core thickness and the pressed application technique

groups. Chi-square revealed a significant effect on mode of fracture for the total thickness and

core thickness (P<0.05), but not for the veneer application method. An increase in the percentage

of the specimens with both core and veneer fracture was observed for the lesser total thickness

group and the greater core thickness group. Significance: The lowest fracture strengths and

greatest incidence of both core and veneer porcelain fracture mode were associated with the 1.5

mm total porcelain thickness groups. A 2 mm thickness of porcelain is preferable in occlusal

stress areas. Pressed veneering porcelain improved the fracture strength of the bilayered zirconia-

porcelain samples in comparison to the manually-applied ones.

Keywords: Fracture load; fracture mode: zirconium oxide ceramics; veneering porcelain;

thickness; thickness ratio

32

Introduction


concerns of metallic restorations stimulated dental manufacturers and dentists to explore

alternatives such as all-porcelain restorations. However, the clinical performance of all-

ceramic restorations has been short of ideal due, to a great extent, to their inherent

brittleness, which makes them more susceptible to fracture, especially in the posterior

region of the mouth.1,2

In the past decade, dental manufacturers focused their efforts on developing new all-

porcelain products with enhanced mechanical properties to improve their clinical

performance in the posterior region through better resistance to chipping and fracture.

Yttrium-tetragonal zirconia poly-crystals (Y-TZP) is one of these products that received

special attention due to its high flexural strength and fracture toughness.3 The tetragonal

crystals in these zirconium oxide ceramics are meta-stable and can be transformed into

larger monoclinic crystals with the application of stress from cracks or flaws.4 This

phenomenon is beneficial in hindering crack growth and increasing fracture toughness;

hence, it is referred to as “transformation toughening”.5 However, Y-TZP-based

products cannot provide optimal esthetics because they have high opacity and are

difficult to stain. Therefore, Y-TZP is frequently used as a core material and coated with

veneering porcelain to enhance esthetics.6

The improved mechanical properties of zirconia-based restorations were not readily

reflected in their clinical performance with several short-term clinical studies reporting a

high incidence of veneering porcelain chipping.7-13

This high chipping rate, in

comparison to that of metal-ceramic restorations, led manufacturers and clinicians to

seek explanations and search for solutions. Several hypothetical explanations have been

reported, including: occlusal overloading, residual stresses as a result of the mismatch in

the coefficient of thermal expansion (CTE) between core and veneering porcelain, 14, 15

residual “tempering” stresses due to the thermal gradient between the veneering

porcelain and the zirconia core,15

interfacial stresses as a result of crystal transformation

at the interface,16

inadequate framework support,17

inherent strength deficiency of the

veneering porcelain 17, 18

and a weak porcelain-zirconia bond.19, 20

Several potential

solutions were suggested in clinical and in-vitro studies to improve the clinical

performance of zirconia-based restorations. These were mainly directed towards

33

improving the strength of the veneering layer and the veneer-to-core bond. The use of the

pressing method to apply the veneering porcelain was suggested as an alternative to the

manual build-up method to decrease the number and size of the internal flaws within the

veneering layer, thus eliminating the potential for micro-crack formation and

propagation.21

Other suggestions were directed towards enhancing the core design to

provide adequate and even support to the veneering porcelain,22, 23

developing stronger

veneering materials9 and modifying the core to veneer thickness ratio to enhance the

overall strength of the crown.24

The aim of this study was to evaluate the effect of varying overall thickness, zirconia

core thickness and veneering porcelain application techniques on the fracture strength of

bilayered zirconia-porcelain discs. Revealing the effects of these factors will improve our

understanding of the strength and limitations of porcelain crowns made with zirconium

oxide cores and will help us in improving the design of the crowns in order to optimize

their strength. The objectives were to test the following null hypotheses: There is no

significant difference in the fracture strength and fracture mode of zirconia-porcelain

specimens fabricated using two different total porcelain thicknesses (1.5 and 2.0 mm),

three different zirconia core thicknesses (0.6, 0.8 and 1.0 mm) and two different

veneering porcelain application techniques (pressed and manual).

34

Materials and Methods

The composition and some of the properties of the materials used (as reported by

manufacturers) are listed in Table 3.1.

Specimen preparation

One hundred and twenty zirconia rectangular specimens (11.5 x 12.5 mm) were

prepared in three thicknesses (0.6, 0.8 and 1.0 mm) (n=40/group). These were sliced

from IPS e.max ZirCAD CEREC blocks (Size B40) (Ivoclar Vivadent, Schaan,

Lichtenstein) with a slow-speed diamond saw (Isomet, Buehler, Lake Buff, IL, USA).

The thickness of the slices was increased by 25% before cutting to compensate for the

shrinkage of zirconia that takes place during sintering. All discs were sintered in a high-

temperature furnace (Sintramat, Ivoclar Vivadent) according to manufacturer’s

instructions. After sintering, the disc thicknesses were measured using a digital calliper

(Mitutoyo Corporation, Tokyo, Japan). A layer of liner (IPS e.max Zirliner, Ivoclar

Vivadent) was applied to one of the surfaces of the disc. The Zirliner powder was mixed

with build-up liquid (Ivoclar Vivadent) to a creamy consistency, then layered onto the

core specimens, vibrated to achieve an even, greenish color, and finally fired in a

porcelain furnace (Whip Mix Pro 100, Whip Mix Corp, Louisville, KY, USA) according

to the manufacturer’s instructions (Table 3.2). The zirconia discs were then veneered

with two veneering porcelains (pressed (P) and manually-applied (M)) on the surface

treated with the IPS e.max Zirliner to total thickness of 1.5 mm or 2.0 mm. Hence, 12

groups of specimens (n=10) were fabricated by varying the total thickness (1.5 mm & 2.0

mm), core thickness (0.6, 0.8, 1.0 mm) and veneer application technique (P, M) (Table

3.3).

Several two-part metallic molds (Figure 3.1) of different thicknesses were used to

apply the veneering porcelain. The molds for the manually applied porcelain were

oversized to accommodate for the anticipated shrinkage (about 20 %). The powder of the

manually applied veneering material (IPS e.max Ceram, Ivoclar Vivadent) was mixed

with the manufacturer’s recommended liquid. The obtained slurry was vibrated and

condensed into the mold and the excess liquid was removed by means of absorbent paper.

The mold was removed and the specimens were fired in a porcelain furnace (Whip Mix

35

Pro 100) according to the manufacturer’s instructions (Table 3.2). For the pressed

veneering specimens, the molds were used to apply a wax pattern on the zirconia discs.

The mold was removed and wax patterns were sprued and invested using a phosphate-

bonded investment material (IPS PressVEST Speed, Ivoclar Vivadent) according to the

manufacturer’s instructions. The ceramic ingot (IPS e.max Zirpress, Ivoclar Vivadent)

was pressed according to the manufacturer’s instructions using a pressing furnace

(EP600 Combi, Ivoclar Vivadent). Firing and subsequent glazing protocols for the

materials used in the study are summarized in Table 3.2.

The veneering porcelain surface was polished with #400, #600, #1000 and #4000

grit silicon carbide paper (Buehler Ltd., Evanston, IL, USA) to a final thickness of 1.5 or

2.0 mm. The disc dimensions were measured using a digital calliper (Mitutoyo

Corporation) (Table 3.3). The specimens were examined under magnification for the

presence of any obvious surface flaws that could affect strength. Any sample with visible

surface imperfections was corrected if possible or discarded. All specimens were glazed

using a glazing paste (IPS e.max Ceram Glaze Paste, Ivoclar Vivadent) which was

applied to the veneering porcelain surface and fired according to the manufacturer’s

instructions (Table 3.2). All specimens were prepared by one operator (M.Z.) in order to

reduce variability.

The discs were cemented to highly filled epoxy resin blocks (5mm in thickness;

Viade Products Inc. Camarillo, CA, USA) which has a modulus of elasticity similar to

human dentin (12.9 GPa) and responds to 34% phosphoric acid-etching to form surface

micro-roughness suitable for bonding.25

First, the under-surface of the zirconia discs was

grit blasted for 5 seconds with 50 μm Al2O3 particles using a micro-etcher under 2 bar

pressure. A bonding/silane coupling agent containing adhesive phosphate monomer

(Clearfil SE bond/Porcelain Bond Activator, Kuraray America Inc., New York, NY,

USA) was applied to the under-surface of the zirconia disc (according to the

manufacturer’s recommendations).26

The epoxy blocks were etched using 40%

phosphoric acid (Kuraray America Inc.) for 1 minute then cleaned using water spray and

dried using oil-free compressed air. All discs were cemented with dual-cured phosphate-

modified resin cement (Panavia F 2, Kuraray America Inc.) according to the

manufacturer’s instructions. One hour after cementation, the specimens were stored in 37

˚C distilled water for 1 week.

36

Fracture load measurement

Each specimen was supported by a flat rigid surface (stainless steel plate). A

tungsten carbide loading ball (8 mm in diameter) was used to load the center of the

veneering porcelain surface (Figure 3.2) in air using a hydraulic-driven universal testing

machine (Instron 8501; Instron, Canton, MA) with a cross-head speed of 1 mm/min.

Fracture was defined as the occurrence of a sharp drop in the load and confirmed with

acoustic events. At the first sign of fracture, the test was terminated. The load-

displacement data from the Instron machine were imported to Microsoft Office Excel and

load-displacement curves were generated. Each curve was analyzed to determine the

failure point (the first sharp drop in the curve) and the corresponding load (N) was

recorded as the failure load.

All fractured specimens were examined under magnification (2.5X) to determine the

mode of fracture (if it involved the veneering porcelain layer alone or both the core and

veneering porcelain layers).

Statistical analyses

Factorial analysis of variances (ANOVA), followed by Bonferroni post-hoc

comparisons, was used to evaluate the effect of veneering material application techniques,

core thickness and total thickness on fracture strength. Chi-square or Fisher’s exact test

was used to evaluate the effect of veneering material application techniques, core

thickness and total thickness on the fracture mode.

All statistical analyses were conducted using the SPSS software for Windows

(version 20, SPSS Inc., IBM, Somers, New York, USA). All statistical analyses were

two-tailed at a significance level of (0.05), except for post hoc analyses of the core

thickness group which used a Bonferroni correction (P=0.05/3=0.017).

37

Results

Fracture load

The loads at fracture of all groups are reported in Table 3.4. Factorial ANOVA

revealed significant main effects of the veneering porcelain application method, core

thickness and total thickness on the fracture load (P ≤0.001) but not for the interaction

terms (Table 3.5). Table 3.6 shows the effect of the independent variables on load at

fracture.

Fracture mode

Distributions of the different fracture modes among the test groups are shown in

Table 3.7. Approximately 58% of the total number of specimens showed a pattern of

fracture that involved the veneering porcelain layer only. Chi-square test revealed

significant effects on the fracture mode of both the core thickness and the total thickness

(P ≤0.001 and P =0.002, respectively), but not the veneering porcelain application

method (Table 3.8). Decreasing the specimen thickness from 2 to 1.5 mm or increasing

the core thickness from 0.6 mm to 1 mm increased the percentage of both core and

veneer fracture.

Representative images of the fractured specimens can be seen in Figure 3.3.

38

Discussion

Most manufacturers are reporting the flexural strength of zirconium oxide cores and

veneering porcelain separately, despite the fact that cores and veneering porcelains are

not separate in the real clinical situation. In this study, composite specimens of zirconia

cores and veneering porcelains were tested in a manner more representative of the

clinical situation. A simple geometrical form of specimens was followed instead of an

actual crown in order to simplify analysis of the variables and to avoid the complexity of

stresses that develop in anatomical crowns. The recommendations for a clinically-

relevant in-vitro load-to-failure test for all-ceramic restorations described by Kelly were

followed in this study,27

including specimen preparation according to the clinical

standards and with clinically-relevant dimensions, and the use of a die material with

elastic modulus similar to that of dentin to avoid variability associated with natural teeth.

The total thicknesses were selected to reflect the recommended reduction at the

functional and non-functional cusps in molars (2 and 1.5 mm, respectively).

In this study, specimens with pressed veneering porcelain had significantly higher

fracture load than the ones with manually-applied porcelain. A similar finding was

previously reported.28

Although both veneering porcelains have a similar fluorapatite

glass-ceramic composition, the flexure strength of the pressed veneering porcelain is

higher than the manually-applied porcelain according to manufacturers’ reports (Table

3.1). The manually-applied technique has the risk of void formation. A void within the

veneering porcelain could act as a flaw for initiation of fracture within the veneering

porcelain and contribute to lower fracture strength results. Because porcelain fracture

resistance is improved by bonding to an underlying more resilient substrate, a more

intimate contact between the zirconia core and veneering porcelain and fewer interfacial

voids in the pressed technique29

, due to the applied pressure under vacuum, likely

enhanced the ability of the pressed veneering porcelain to withstand the applied forces.

The higher fracture loads for specimens with 2 mm total thickness in comparison to

1.5 mm total thickness might be related to the increase in volume of the specimen. It has

to be emphasized that the increase in the total thickness (d) for samples with similar core

thickness resulted from increasing the thickness of veneering layer (dv), which increases

the critical load for radial crack in the core and cone crack in the veneer layer. This

increase is higher for the radial cracking relative to the cone cracking, which renders the

39

specimen more susceptible to cone cracking of the veneering layer 30

and might explain

the increase in the fracture loads and the tendency towards more fractures involving the

veneering layer only by the 2 mm total thickness group.

Significant effects of the core thickness on the fracture strength and fracture mode

were found in this study. Higher fracture loads and lower percentages of fractures

involving the veneer layer only were observed with increasing the core thickness.

Although there was no significant interaction effect, it appears that the fracture strength

of the 1.5mm total thickness specimens was more affected by a reduced core thickness

(and hence, a reduced core to veneer thickness ratio) than the 2mm total thickness

specimen (Figure 3.4). Other studies have reported an increase in strength with the

increase in the core to veneer thickness ratio.6, 31, 32

White et al. used (4x4x40 mm) beams,

representing fixed dental prosthesis but not dental crowns, and reported that increasing

the core to veneer thickness ratio increased the modulus of rupture of beams that were

composed of zirconia core and matching veneering porcelain materials using 3-point

bending test.6 Fleming et al. used a biaxial flexural test to assess the effect of varying the

aluminous core: porcelain thickness ratio and found similar results.31

An increase in the

number of fracture fragments, Hertzian cone and delamination were associated with a

decrease in core to veneer thickness ratio. Wakabayashi and Anusavice32

used bilayered

alumina-porcelain discs supported with different substrates on a design similar to the one

used in this study. They reported that increasing the core to veneer thickness ratio

increased the flexure strength and shifted the crack initiation site from veneer to core.

However, the core to veneer thickness ratios varied in two of the above studies within a

fixed overall thickness 6, 32

, while in the third study,31

the ratio was varied by changing

the core, veneer and overall thickness. It must be emphasized that comparing the results

of the present study with those of the above studies might be difficult due to variability in

materials, thicknesses, specimen designs and test methods.

There appeared to be an increasing incidence of both core and veneer fractures in the

1.5 mm total thickness specimens when the core thickness increased from 0.6 to 1mm,

and the incidence of those fractures was near 100% when the core thickness was 1 mm.

The change in the fracture mode with the increase in the core thickness for specimens

with 1.5 mm overall thickness can be explained by the shielding effect of the

compressive stresses under the loading ball within the thin veneering layer which acts to

40

protect the veneer from developing cone cracks, resulting in increasing the combined

failures of both veneer and core. This shielding effect faded in specimens with increased

veneer thickness, with subsequent increase in the percentages of failures involving the

veneering porcelain only. In contrast, varying the core thickness did not appear to affect

the fracture mode among specimens with 2 mm total thickness. This supports fracture

load and mode insensitivity to variation in the core to veneer thickness ratio among the

specimens with veneering porcelain thickness > 1 mm and reflects the importance of the

veneering porcelain thickness as an influential factor on the fracture load and mode of the

zirconia-porcelain specimens (Figure 3.4 and Table 3.7). It is clear that the impact of the

fracture within the veneering layer only might not be as serious as the bulk fracture. The

latter will require the replacement of the restoration, while the earlier might be addressed

by re-contouring or chairside repair.

The high chipping rate of the veneering porcelain on zirconia framework triggered

the researchers to study the stress profile in an attempt to correlate the stresses within the

veneering layer with the chipping behavior. Residual stresses within the veneering layer

can be generated as a result of thermal gradients occurring during the cooling of the

veneer during fabrication and the mismatch in thermal expansion properties between the

core and veneering porcelain.33

The thermal gradient within the bilayered system is

determined by the thermal properties of both layers, thickness and cooling rate. Mainjot

et al. in a series of studies evaluated the effect of cooling rate and framework and veneer

thickness on the stress profile within the veneering porcelain using the hole-drilling

method.33-36

In one study, zirconia discs (thickness from 0.5 to 3 mm) were veneered

with 1.5 mm veneering porcelain. Specimens with 1.5 mm zirconia had the most

favorable stress profile with internal compressive stresses near the veneer-core interface,

while all other framework thicknesses exhibited either surface or internal tensile

stresses.35

In another study,36

1 mm thick zirconia discs were veneered using different

thicknesses ranging from 1 to 3 mm. All specimens, except the ones with 1.5 mm veneer,

exhibited compressive stresses at the surface of the veneer that changed to tensile nature

at the 0.5 -1 mm depth and then reverted to compressive near the framework. In contrast,

specimens with 1.5 mm veneer exhibited only interior tensile stresses. These studies

varied the core to veneer thickness ratio by varying the thicknesses of the core or the

veneer without fixing the total thickness, which makes it difficult to separate the effect of

41

the ratio from that of the total thickness and to correlate the results of these studies with

the findings of the present study. Perhaps future work should be directed towards

correlating the residual stresses within the veneering porcelain of zirconia-porcelain

bilayered specimens with fracture strength and mode.

Since all loading in this study was uniaxial and other forces were ignored due to the

machine limitations, the clinical implication of the results of the present study should be

interpreted with caution. Fracture loads were recorded when there was a sharp drop in the

load-displacement curve accompanied by acoustic sound. However, the load at complete

fracture may not necessarily be the load that the material will withstand before crack

initiation. Generally, crack initiation occurs at a lower force compared to the one of

complete fracture. Once a crack is initiated, crack propagation starts leading to complete

fracture. This study did not record the “pop-in” crack (crack initiation) of the specimens

because the crack initiation was not accompanied by any drop in the load since the

loading ball was entirely supported by intact specimen surface. The use of tungsten

carbide loading ball increased the contact pressure in the specimen compared to the

clinical situation as the contact pressure is influenced by the ratio of the elastic modulus

of the porcelain to that of the loading ball, and by the radius of the loading ball.37

Alternative loading methods could have included the use of a ball with a modulus of

elasticity lower than that of tungsten carbide, a tin sheet between the load applicator and

crown as stress breaker or a stainless steel loading piston with its end machined to a

curvature equivalent to 40-50 mm diameter to reproduce clinically-relevant contact

pressure. However, this increase in the loading stresses will not eliminate the comparison

between the different test groups since all specimens were loaded using similar contact

pressure. Fracture load data reported in this study should be considered as relative, not

“absolute”, values and extrapolation of these in-vitro strength data to clinical

performance must be considered cautiously and within the limitations of the study.2

42

Conclusions

The current study highlighted the effects of the total thickness, core thickness and

veneering porcelain application method on the fracture strength and mode of the

bilayered zirconia/porcelain samples. Using pressed veneering porcelain increased the

fracture strength of the zirconia/porcelain samples compared to manually-applied

veneering porcelain, but had no effect on the fracture mode. The variable effect of the

core thickness among two total thicknesses (1.5 and 2 mm) underlined the importance of

the veneering thickness in influencing fracture strength and mode. Fracture load and

mode of specimens with veneering thickness less than 1 mm were affected by varying the

core thickness. This effect disappeared once the veneering porcelain thickness increased

beyond 1 mm.

43

Acknowledgements:

The current study is part of a PhD thesis submitted to the Faculty of Dentistry,

University of Toronto.

This work was supported by grants from the Faculty of Dentistry Research

Committee, University of Toronto; and Education and Research Foundation of

Prosthodontics. Materials donations were received from Ivoclar-Vivadent and Kuraray;

and technical support provided by Shaw Dental Labs (Toronto) to whom the authors are

grateful.

44

(Table 3.1) Composition and some properties of materials used for core and veneering porcelain.

(as reported by manufacturers)

Material Manufacturer Composition Manufacturing

Technique

CTE

(10-6 .K-1)

Flexure

strength

(MPa)

IPS e.max ZirCad

(Lot: L13763)

Ivoclar Vivadent,

Schaan, Lichtenstein

Yttrium-stabilized

zirconium oxide CAD-CAM

10.75 ( 0.25)

(100 -400 C) 900 (50)

IPS e.max Zirliner

(Lot: L29047)

Ivoclar Vivadent,


Fluorapatite

glass-ceramic

Manual

application - -

IPS e.max Zirpress

(Lot: H25305)

Ivoclar Vivadent,


Fluorapatite

glass-ceramic Pressing

9.75 (0.25)

(100 - 400

C)

110 (10)

IPS e.max Ceram

(Lot: K13416)

Ivoclar Vivadent,


Fluorapatite

glass-ceramic

Manual

application

9.5 (0.25)

(100 -400 C) 90 (10)

45

(Table 3.2) Firing program for materials used in the study.

Material Pre-drying

temperature

(ºC)

Holding time

(min.)

Heating rate

(ºC/min)

Final firing

temperature

(ºC)

Holding time

(min)

Vacuum

1

(ºC)

Vacuum

2

(ºC)

IPS e.max Zirliner 403 4 40 960 1 450 959

IPS e.max Ceram 403 4 40 750 1 450 749

IPS e.max ZirPress (200g

ring) 700 - 60 910 15 500 910

Glazing 403 6 60 725 1 450 724

46

(Table 3.3) Experimental groups according to core thickness (mm), total thickness (mm) and veneering

porcelain.

Group

Core thickness

Mean (SD)

Total thickness

Mean (SD) Veneering Porcelain

0.6/1.5 P 0.57 (0.02) 1.47 (0.03) Pressed

0.8/1.5 P 0.77 (0.02) 1.49 (0.02) Pressed

1/1.5 P 0.99 (0.01) 1.50 (0.01) Pressed

0.6/1.5 M 0.59 (0.02) 1.49 (0.03) Manually-applied

0.8/1.5 M 0.79 (0.02) 1.50 (0.01) Manually-applied

1/1.5 M 0.99 (0.02) 1.52 (0.03) Manually-applied

0.6/2 P 0.59 (0.03) 1.97 (0.04) Pressed

0.8/2 P 0.78 (0.02) 1.99 (0.04) Pressed

1/2 P 0.99 (0.01) 2.01 (0.03) Pressed

0.6/2 M 0.60 (0.02) 2.00 (0.04) Manually-applied

0.8/2 M 0.79 (0.03) 2.01 (0.04) Manually-applied

1/2 M 0.99 (0.01) 2.00 (0.02) Manually-applied

47

(Table 3.4) Fracture load (N) data for all groups.

Total

Thickness

Core

Thickness

Pressed Manually-applied

Mean (SD)

95 %

Confidence

Interval

Min - Max Mean (SD)

95 %

Confidence

Interval

Min - Max

1.5 mm

0.6 mm 1145 (199) 1003 – 1288 697 – 1335 780 (262) 610 - 985 456 - 1227

0.8 mm 1300 (183) 1169 – 1431 1051 – 1635 1119 (166) 1000 - 1237 762 - 1302

1 mm 1347 (182) 1217 – 1477 1122 – 1719 1361 (182) 1231 – 1491 1151 – 1680

2 mm

0.6 mm 1499 (394) 1217 – 1781 599 – 1913 1341 (248) 1163 – 1517 858 - 1621

0.8 mm 1465 (306) 1246 – 1683 943 – 1908 1444 (374) 1176 - 1711 1035 - 2147

1 mm 1465 (306) 1246 – 1683 943 – 1908 1363 (318) 1135 – 1590 854 - 1849

48

(Table 3.5) Factorial ANOVA

(dependent variable: fracture load; independent variables: veneering porcelain application method, core

thickness and total thickness).

R2 = .407 (Adjusted R2 = .347)

Source Type III Sum of

Squares df Mean Square F Sig.

Corrected Model 5415777.056 11 492343.369 6.742 ≤0.001

Intercept 209468718.591 1 209468718.591 2868.514 ≤0.001

Core Thickness 1160804.299 2 580402.150 7.948 ≤0.001

Total Thickness 2445499.931 1 2445499.931 33.489 ≤0.001

Veneer Material 841776.947 1 841776.947 11.527 ≤0.001

Core Thickness * Total Thickness 430515.273 2 215257.636 2.948 0.057

Core Thickness * Veneer Material 121024.517 2 60512.259 0.829 0.439

Total Thickness * Veneer Material 538.624 1 538.624 0.007 0.932

Core Thickness * Total Thickness *

Veneer Material 415617.464 2 207808.732 2.846 0.062

Error 7886529.228 108 73023.419

Total 222771024.875 120

Corrected Total 13302306.284 119

49

(Table 3.6) The effect of different independent variables on fracture load (N).

* Identical letters indicate no statistical difference based on Bonferroni’s test (p > 0.0167).

Independent Variable Mean (SD) 95 % Confidence

Interval

Veneering Porcelain

Application

Technique

Pressed 1405 (310) 1325 – 1485

Manually-applied 1237 (339) 1184 – 1365

Total Thickness

1.5mm 1178 (273) 1129 – 1269

2mm 1464 (331) 1397 – 1566

Core Thickness *

0.6 mm A 1196 (381) 1112 – 1365

0.8 mm A, B 1332 (296) 1248 – 1442

1 mm B 1436 (280) 1346 – 1526

50

(Table 3.7) Incidence of different types of fracture mode for all groups.

Total

Thickness

Core

Thickness

Veneer

Thickness

Pressed Manually-applied

Veneer Only

N (%)

Core and

Veneer

N (%)

Veneer Only

N (%)

Core and

Veneer

N (%)

1.5 mm

0.6 mm 0.9 mm 6 (60 %) 4 (40 %) 8 (80 %) 2 (20 %)

0.8 mm 0.7 mm 5 (50 %) 5 (50 %) 5 (50 %) 5 (50 %)

1 mm 0.5 mm 0 10 (100 %) 1 (10 %) 9 (90 %)

2 mm

0.6 mm 1.4 mm 8 (80 %) 2 (20 %) 8 (80 %) 2 (20 %)

0.8 mm 1.2 mm 8 (80 %) 2 (20 %) 7 (70 %) 3 (30 %)

1 mm 1 mm 7 (70 %) 3 (30 %) 7 (70 %) 3 (30 %)

51

(Table 3.8) Effect of different independent variables on fracture mode

Independent Variable Veneer Only

N (%)

Core and Veneer

N (%)

Significance

Level *

Veneering Porcelain

Application Technique

Pressed 34 (56.7 %) 26 (43.3 %)

P=0.711

Manually applied 36 (60%) 24 (40 %)

Total Thickness

1.5mm 25 (41.7 %) 35 (58.3 %)

P<0.001

2mm 45 (75 %) 15 (25 %)

Core Thickness

0.6 mm 30 (75 %) 10 (25%)

P= 0.002 0.8 mm 25 (62.5 %) 15 (37.5 %)

1 mm 15 (37.5 %) 25 (62.5 %)

* using Chi-Square test.

52

(Figure 3.1) Two of the two-piece split molds used to fabricate the specimens.

(Figure 3.2) Porcelain- Zirconia Specimen

a) Top and cross section views

b) Specimen loading (cross section)

53

(Figure 3.3) Examples of the fractured samples

a) Specimen from group 0.8/2 M showing a cohesive fracture within the veneering porcelain

b) Specimen from group 0.6/2 P showing combined fracture (adhesive (hollow arrow) and cohesive (solid

arrow)) involving the veneering porcelain.

c) Specimen from group 1/1.5 M showing a fracture involving both the core and veneering porcelain. Note

the crack (arrow) that penetrated all the way through the core layer.

d) Specimen from group 0.8/1.5 P showing a fracture involving both the core and veneering porcelain. Note

that the crack (arrow) did not penetrate through the core layer, which indicates that the crack started at the

core-cement interface.

54

(Figure 3.4) Error bar graph for the effect of core thickness and veneering application method on the fracture

load.

a) 1.5 mm and b) 2 mm total-thickness samples.

B

A

55

References

1. Goodacre CJ, Bernal G, Rungcharassaeng K, Kan JY. Clinical complications in

fixed prosthodontics. J Prosthet Dent 2003;90(1):31-41.

2. Kelly JR. Dental ceramics: current thinking and trends. Dent Clin North Am

2004;48(2):viii, 513-viii, 30.

3. Denry I, Kelly JR. State of the art of zirconia for dental applications. Dent Mater

2008;24(3):299-307.

4. Albakry M, Guazzato M, Swain MV. Effect of sandblasting, grinding, polishing

and glazing on the flexural strength of two pressable all-ceramic dental materials.

J Dent 2004;32(2):91-9.

5. Christel P, Meunier A, Heller M, Torre JP, Peille CN. Mechanical properties and

short-term in-vivo evaluation of yttrium-oxide-partially-stabilized zirconia. J

Biomed Mater Res 1989;23(1):45-61.

6. White SN, Miklus VG, McLaren EA, Lang LA, Caputo AA. Flexural strength of

a layered zirconia and porcelain dental all-ceramic system. J Prosthet Dent

2005;94(2):125-31.

7. Raigrodski AJ, Hillstead MB, Meng GK, Chung KH. Survival and complications

of zirconia-based fixed dental prostheses: a systematic review. J Prosthet Dent

2012;107(3):170-7.

8. Cehreli MC, Kokat AM, Akca K. CAD/CAM Zirconia vs. slip-cast glass-

infiltrated Alumina/Zirconia all-ceramic crowns: 2-year results of a randomized

controlled clinical trial. J Appl Oral Sci 2009;17(1):49-55.

9. Heintze SD, Rousson V. Survival of zirconia- and metal-supported fixed dental

prostheses: a systematic review. Int J Prosthodont 2010;23(6):493-502.

10. Ortorp A, Kihl ML, Carlsson GE. A 3-year retrospective and clinical follow-up

study of zirconia single crowns performed in a private practice. J Dent

2009;37(9):731-6.

11. Hosseini M, Worsaae N, Schiodt M, Gotfredsen K. A 1-year randomised

controlled trial comparing zirconia versus metal-ceramic implant supported

single-tooth restorations. Eur J Oral Implantol 2011;4(4):347-61.

12. Rinke S, Schafer S, Roediger M. Complication rate of molar crowns: a practice-

based clinical evaluation. Int J Comput Dent 2011;14(3):203-18.

13. Sagirkaya E, Arikan S, Sadik B, Kara C, Karasoy D, Cehreli M. A randomized,

prospective, open-ended clinical trial of zirconia fixed partial dentures on teeth

and implants: interim results. Int J Prosthodont 2012;25(3):221-31.

14. de Kler M, de Jager N, Meegdes M, van der Zel JM. Influence of thermal

expansion mismatch and fatigue loading on phase changes in porcelain veneered

Y-TZP zirconia discs. J Oral Rehabil 2007;34(11):841-7.

15. Swain MV. Unstable cracking (chipping) of veneering porcelain on all-ceramic

dental crowns and fixed partial dentures. Acta Biomater 2009;5(5):1668-77.

56

16. Tholey MJ, Swain MV, Thiel N. SEM observations of porcelain Y-TZP interface.

Dent Mater 2009;25(7):857-62.

17. Sailer I, Feher A, Filser F, Gauckler LJ, Luthy H, Hammerle CH. Five-year

clinical results of zirconia frameworks for posterior fixed partial dentures. Int J

Prosthodont 2007;20(4):383-8.

18. Raigrodski AJ, Chiche GJ, Potiket N, Hochstedler JL, Mohamed SE, Billiot S, et

al. The efficacy of posterior three-unit zirconium-oxide-based ceramic fixed

partial dental prostheses: a prospective clinical pilot study. J Prosthet Dent

2006;96(4):237-44.

19. Aboushelib MN, de Jager N, Kleverlaan CJ, Feilzer AJ. Microtensile bond

strength of different components of core veneered all-ceramic restorations. Dent

Mater 2005;21(10):984-91.

20. Guess PC, Kulis A, Witkowski S, Wolkewitz M, Zhang Y, Strub JR. Shear bond

strengths between different zirconia cores and veneering ceramics and their

susceptibility to thermocycling. Dent Mater 2008;24(11):1556-67.

21. Aboushelib MN, de Kler M, van der Zel JM, Feilzer AJ. Effect of veneering

method on the fracture and bond strength of bilayered zirconia restorations. Int J

Prosthodont 2008;21(3):237-40.

22. Bonfante EA, da Silva NR, Coelho PG, Bayardo-Gonzalez DE, Thompson VP,

Bonfante G. Effect of framework design on crown failure. Eur J Oral Sci

2009;117(2):194-9.

23. Rosentritt M, Steiger D, Behr M, Handel G, Kolbeck C. Influence of substructure

design and spacer settings on the in vitro performance of molar zirconia crowns. J

Dent 2009;37(12):978-83.

24. Zarone F, Russo S, Sorrentino R. From porcelain-fused-to-metal to zirconia:

clinical and experimental considerations. Dent Mater 2011;27(1):83-96.

25. Neiva G, Yaman P, Dennison JB, Razzoog ME, Lang BR. Resistance to fracture

of three all-ceramic systems. J Esthet Dent 1998;10(2):60-6.

26. Blatz MB. Cementation of zirconium-oxide ceramic restorations. Pract Proced

Aesthet Dent 2004;16(1):14.

27. Kelly JR. Clinically relevant approach to failure testing of all-ceramic

restorations. J Prosthet Dent 1999;81(6):652-61.

28. Lin WS, Ercoli C, Feng C, Morton D. The effect of core material, veneering

porcelain, and fabrication technique on the biaxial flexural strength and weibull

analysis of selected dental ceramics. J Prosthodont 2012;21(5):353-62.

29. Aboushelib MN, Kleverlaan CJ, Feilzer AJ. Microtensile bond strength of

different components of core veneered all-ceramic restorations. Part II: Zirconia

veneering ceramics. Dent Mater 2006;22(9):857-63.

30. Bhowmick S, Melendez-Martinez JJ, Zhang Y, Lawn BR. Design maps for

failure of all-ceramic layer structures in concentrated cyclic loading. Acta Mater

2007;55(7):2479-88.

57

31. Fleming GJ, El-Lakwah SF, Harris JJ, Marquis PM. The effect of core:dentin

thickness ratio on the bi-axial flexure strength and fracture mode and origin of

bilayered dental ceramic composites. Dent Mater 2005;21(2):164-71.

32. Wakabayashi N, Anusavice KJ. Crack initiation modes in bilayered

alumina/porcelain disks as a function of core/veneer thickness ratio and

supporting substrate stiffness. J Dent Res 2000;79(6):1398-404.

33. Mainjot AK, Schajer GS, Vanheusden AJ, Sadoun MJ. Influence of cooling rate

on residual stress profile in veneering ceramic: measurement by hole-drilling.

Dent Mater 2011;27(9):906-14.

34. Mainjot AK, Schajer GS, Vanheusden AJ, Sadoun MJ. Residual stress

measurement in veneering ceramic by hole-drilling. Dent Mater 2011;27(5):439-

44.

35. Mainjot AK, Schajer GS, Vanheusden AJ, Sadoun MJ. Influence of zirconia

framework thickness on residual stress profile in veneering ceramic:

measurement by hole-drilling. Dent Mater 2012;28(4):378-84.

36. Mainjot AK, Schajer GS, Vanheusden AJ, Sadoun MJ. Influence of veneer

thickness on residual stress profile in veneering ceramic: measurement by hole-

drilling. Dent Mater 2012;28(2):160-7.

37. Lawn BR. Fracture of brittle solids. Cambridge solid state science series. 2nd ed.

Cambridge ; New York: Cambridge University Press; 1993. p. 249-306.

58


59

Effect of core thickness and design on fracture strength and fatigue

resistance of zirconia-based crowns

Authors:




Saudi Arabia


Associate Professor Dentistry, Schulich School of Medicine & Dentistry, The University of

Western Ontario, London, ON, Canada






Toronto, ON, Canada


Laura E.Tam, DDS, MSc


Toronto, ON, Canada






Phone: (416) 979-4934 X 4572

Fax: (416) 979-4936


60

Abstract

Purpose: to evaluate the influence of varying the core thickness and modifying the core

design on the fracture strength, mode and extent of zirconia-based crowns and to compare it

to the metal-ceramic crowns. Methods: An ivorine molar was prepared to receive an all-

ceramic crown. Using an epoxy resin, 50 replication dies were made of the prepared tooth.

Forty all-ceramic crowns were made using zirconium-oxide cores (IPS e.max ZirCAD) and

pressed veneering porcelain (IPS e.max Zirpress). All-ceramic crowns were fabricated

according to four different coping designs (n=10/group): 0.6mm even-thickness (group 1),

0.6mm anatomically-shaped (group 2), 1mm even-thickness (group 3), and 1mm

anatomically-shaped (group 4). In addition, ten metal-ceramic crowns were fabricated as

controls. All crowns were cemented to their respective dies using resin cement (PanaviaF

2.0). After water storage, crowns were subjected to cyclic loading in a universal testing

machine (50–600N for 500,000 cycles at 20Hz). After fatigue, crowns were loaded to

fracture at crosshead speed of 0.5 mm/min. Fracture load, mode and extent were recorded.

Data were analyzed using factorial ANOVA and chi-square or Fisher’s exact test at α=0.05.

Results: All the all-ceramic crowns survived the fatigue test, while two metal-ceramic

crowns failed during the fatigue testing. Factorial ANOVA revealed a statistically significant

effect for the core thickness (Mean [SD]: 0.6 mm= 2,543 [716] N and 1mm= 3,061 [816] N),

but not for the core design and the interaction term. Metal-ceramic crowns showed

significantly higher mean fracture load in comparison to group1. All the all-ceramic crowns

except two showed fractures involving the veneering porcelain only. Similarly, all the metal-

ceramic crowns showed fractures involving the veneering layer only with exposure of the

metal copings. Fisher’s exact test revealed no significant effect for the core thickness and

design on the fracture mode. In contrast, core design but not thickness influenced the fracture

extent. Higher percentages of failure involving the axial walls were observed for the even

thickness copings. Conclusions: The core thickness, but not the core design, had a

significant effect on fracture strength of zirconia-based crowns. Anatomically shaped

copings reduced the extent of the veneering porcelain fracture.

Keywords: Fracture strength; fracture mode; zirconium oxide; Zirconia core design;

thickness; crown; ceramics

61

Introduction



alternatives such as all-porcelain restorations. However, the clinical performance of all-

ceramic restorations has been short of ideal due, to a great extent, to their inherent



In a systematic review on clinical complications in fixed prosthodontics,1 all-ceramic

crowns showed an 8% incidence of complications, with crown fractures being the most

common. Molars showed a higher fracture rate (21%) than premolars and anterior teeth

(7% and 3%, respectively). Kelly reported similar results with higher fracture rates in

posterior crowns compared to anterior crowns.2 This higher fracture rate in the posterior

region can be explained by the higher biting forces in the posterior region (400 to 890 N

in the molar region and 222 to 445 N in the premolar area) in comparison to the anterior

region (133 to 334 N in the cuspid area and 89 to 111 N in the incisor region).3

In the past decade, dental manufacturers have focused their efforts on developing

new all-porcelain products with enhanced mechanical properties to improve their clinical

performance in the posterior region through better resistance to chipping and fracture.

Yttrium-tetragonal zirconia poly-crystals (Y-TZP) is one of these products that has

received special attention due to its high flexural strength and fracture toughness.4 The

tetragonal crystals in these zirconium oxide ceramics are meta-stable and can be

transformed into larger monoclinic crystals with the application of stress from cracks or

flaws.5 This phenomenon is beneficial in hindering crack growth and increasing fracture

toughness; hence, it is referred to as “transformation toughening”.6 However, Y-TZP-

based products cannot provide optimal esthetics because they have high opacity and are

difficult to stain. Therefore, Y-TZP must be used as a core material and clad with

veneering porcelain to enhance esthetics. 7

The improved mechanical properties of zirconia-based restorations were not readily

reflected in their clinical performance with several short-term clinical studies reporting a

high incidence of veneering porcelain chipping with rare framework fractures.8-14

62

Although minor veneering porcelain chipping, in most of the cases, does not necessitate

replacement of the restoration, it may create rough surface or sharp edge causing patient

discomfort or if involving the esthetic zone, it may compromise the esthetics. This high

chipping rate, in comparison to that of metal-ceramic restorations, led manufacturers and

clinicians to seek explanations and search for solutions. Several hypothetical

explanations have been reported, including: occlusal overloading, residual stresses as a

result of the mismatch in the coefficient of thermal expansion (CTE) between core and

veneering porcelain, 15, 16

residual “tempering” stresses due to the thermal gradient

between the veneering porcelain and the zirconia core,16

interfacial stresses as a result of

crystal transformation at the interface,17


inherent

strength deficiency of the veneering porcelain 18, 19

and a weak porcelain-zirconia bond.20,

21 Several potential solutions were suggested in clinical and in-vitro studies to improve

the clinical performance of zirconia-based restorations. These were mainly directed

towards improving the strength of the veneering layer and the veneer-to-core bond. Two

of these suggestions were directed towards enhancing the core design to provide

adequate and even support to the veneering porcelain,22

and modifying the core:veneer

thickness ratio to enhance the overall strength of the crown.23

At the beginning of their introduction, zirconia copings were fabricated with uniform

thickness. As a consequence, the veneering porcelain thickness varied at different regions

of the restoration with the veneering porcelain tending to be thick in some areas,

especially in the cusp areas. It has to be highlighted that the fracture toughness of most

veneering porcelains is about eight times lower than that of the zirconia core. 24

Hence, it

can be expected that restorations with thick veneering porcelain will have a lower load

bearing capacity than those with thinner veneering layer. 25

The thickness of the

veneering layer can be reduced by modifying the coping design to provide even veneer

thickness and/or by increasing the core thickness. At the early stages of the zirconia-

based restorations, the modification of the coping design was time consuming and

required a cut back of a full-contoured restoration due to limitation in the designing

software. Currently, the evolution of CAD/CAM systems allows the virtual designing of

an “anatomic” core.

63

Several in-vitro studies evaluated the effect of core design modification on the

performance of zirconia-based restorations. 25-29

These studies used different

methodologies and provided conflicting results. Clinical studies on zirconia fixed partial

dental prosthesis with anatomic framework design showed promising results, but these

studies were short-term case series without controls.

When zirconia-based ceramic were first introduced to dentistry, manufacturers

recommended a minimal thickness of 0.6-0.7 mm to provide adequate strength to the

core without any recommendation for the core to veneer thickness ratio. A high core to

veneer thickness ratio was shown to improve the in-vitro fracture resistance of the flat

zirconia-porcelain bilayer samples.7, 30

However, the effect of the core to veneer

thickness ratio on the in-vitro or in-vivo performance of the zirconia-based crowns has

not been as yet adequately-addressed in the literature.

Thus, the aim of this study was to evaluate the influence of varying the core

thickness and modifying the core design on the fracture strength and mode of zirconia-

based crowns and to compare it to that of the metal-ceramic crowns. The objectives were

to test the following null hypotheses: 1) There is no significant difference in the fracture

strength, mode and extent of zirconia-based crowns fabricated using two different core

thickness (0.6 and 1 mm) and two different zirconia core designs (even thickness and

anatomically shaped), and 2) there is no significant difference in the fracture strength and

fracture mode between zirconia-based and metal-ceramic crowns.

64

Materials and Methods

An ivorine mandibular molar replica was prepared to receive an all-ceramic crown

following dimensions recommended for the CEREC system (Sirona Dental Systems

GmbH, Bensheim, Germany) (Figure 4.1). The tooth had an occlusal reduction of 1.5

mm and minimum axial reduction of 1 mm with a 6 to 8° total angle of convergence (3 to

4° on each side). The gingival margin was prepared with a circumferential shoulder at

least 1 mm wide. Lingual and facial surfaces were prepared in two planes, and all line

angles were smoothed to reduce stress concentration. The prepared ivorine molar was

used as a master die to fabricate 50 replicas using a highly filled epoxy resin (Viade

Products, Inc., Camarillo, CA) with a modulus of elasticity similar to that of human

dentin (12.9 GPa). This material responded to 34% phosphoric acid etching by forming a

surface microroughness suitable for bonding. 31

Coping design and fabrication

The prepared molar replica was inserted in a dentiform with mesial and distal

adjacent ivorine teeth in place. The surfaces were then covered with a thin layer of

optical reflective medium (CEREC Optispray, Sirona Dental Systems GmbH). A scan of

the prepared tooth with its adjacent teeth was then captured using the inEos Blue scanner

(Sirona Dental Systems GmbH). A crown was designed using the biogeneric mode of the

inLab software (Version 3.85, Sirona Dental Systems GmbH), which had a minimum

thickness of 1.5 mm at the central fossa and 2mm at the cusp area, and a reference master

crown was milled from Vita Mark II blocks (VitaZahnfabrik, BadSackingen, Germany)

using the inLab MCXL milling unit (Sirona Dental Systems GmbH).

Four different coping designs (Figure 4.2a) were made using the inLab software

(Version 3.85) as follows:

Non-anatomical copings with even thickness of 0.6 mm occlusally and

circumferentially (group 1).

Anatomically shaped copings with a minimum thickness of 0.6 mm

occlusally and circumferentially (group 2).

65

Non-anatomical copings with even thickness of 1 mm occlusally and 0.6 mm


Anatomically shaped copings with a minimum occlusal thickness of 1 mm

and 0.6 mm circumferential thickness (Group 4).

The non-anatomical copings were designed using the framework mode, while the

anatomical ones were designed to keep a constant thickness of the veneering porcelain

using the reduction mode and the master crown as a reference for the occlusal anatomy.

The thickness of the master crown was reduced by 0.9 mm and 0.5 mm to produce

copings with minimum thickness of 0.6 mm and 1 mm, respectively. All copings were

designed with a 0.5 mm circumferential collar. Forty copings were milled (10/group)

from IPS e.max ZirCAD blocks (size C15) (Ivoclar Vivadent, Schaan, Liechtenstein)

using the inLab MCXL milling unit (Sirona Dental Systems GmbH). All copings were

dried and sintered according to the manufacturer’s instructions. The sintering procedure

was conducted in a sintering furnace (Zyrcomat T furnace, Vita Zahnfabrik) at a

temperature of 1530◦C. The duration of the program, including the cooling phase, was

approximately 8 hours.

Veneering layer fabrication.

First, a thin layer of liner material (IPS e.max Ceram ZirLiner, Ivoclar Vivadent)

was applied on all the copings and fired according to the manufacturer’s instructions. A

coping from each group was then scanned using the inEos Blue scanner and a veneering

layer was designed with the inLab software using the master crown as a reference in the

correlation mode. The veneering layer designs were exported to the inLab Stack software

(Sirona Dental Systems GmbH), which allows the processing of several restorations in

one milling procedure using a single block. Special acrylic polymer blocks (Size CW-40)

(VITA CAD-Waxx, VitaZahnfabrik), which burn without residue, were used to mill the

veneering layers (3 veneering structures/block) using the inLab MCXL milling unit.

After milling, the veneering structures were fitted on their respective copings and the

junction between the coping and veneering structure was sealed using casting wax

(Figure 4.2b). The CAD-waxx veneering structures were smoothed, sprued and invested

66

into a special investing material (Galaxy II, Talladium Inc., Valencia, CA, USA) in a

200g-investing ring (Ivoclar Vivadent) (3 crowns/ring) according to the manufacturer’s

instructions. The investment ring was placed in a burnout oven (Infinity L30, Jelrus

International, Melville, NY, USA) and heated at 1562 °C for 45 minutes, then directly

transferred to the ceramic press oven. The copings were overpressed by a fluorapatite

glass-ceramic material (IPS e.max ZirPress, Ivoclar Vivadent). After cooling to room

temperature, the investment was divested and immersed into a hydrofluoric acid solution

(IPS e.max Press Invex Liquid, Ivoclar Vivadent) for 5 min in an ultrasonic bath to

remove the reaction layer. Afterward, the crowns were cleaned under running water for

3min and dried. The pressing sprues and extrusion flashes were removed using a water-

cooled air-turbine without pressure to protect the porcelain from heat damage. Finally,

the crowns were polished with a porcelain polishing kit (Dialite extra-oral porcelain

polishing kit, Brasseler USA, Savannah, GA, USA) (Figure 4.2c).

Metal-ceramic crown fabrication (Control Group).

Ten metal-ceramic crowns were fabricated as a control group. First, coping replicas

were milled from CAD-waxx blocks (Size CW-40) using the same data file that was used

to mill the copings in Group 2 (4.2d). Then, the coping replica was sprued, invested into

a phosphate-bonded investment (Bellavest SH, Bego, Bremen, Germany) and cast using

non-precious alloy for metal-ceramic restorations (Authentic Alloy HE; Ceramay,

Stuttgart, Germany) according to the manufacturer’s recommendations (Figures 4.2e and

4.2f-1). Similar to the zirconia crowns, the veneering layers were milled from CAD-

waxx blocks (Size CW-40) using the data file that was used to fabricate the veneering

layer in Group 2 (Figure 4.2f-2). The veneering replicas were fitted and sealed on the

metal copings (Figure 4.2f-3), sprued and invested using a universal investment (Galaxy

II). The wax was burnt out and the investment ring was preheated according to the

manufacturer’s instructions. Ceramic ingot (Authentic, Ceramay) was pressed over the

metal copings according to the manufacturer’s instructions (Figure 4.2g). All the metal-

ceramic crowns were produced by one experienced technician in a commercial dental lab,

while all the zirconia-based crowns were fabricated by the primary author.

67

The details about the materials used in this study and their firing protocol are listed

in Tables 4.1 and 4.2, respectively.

Cementation

Before cementation, the crown thickness at the central fossa was verified to have a

minimum thickness of 1.5 mm using a caliper (Buffalo Dental Manufacturing Co.,

Syosset, NY, USA) and the crown was seated on its respective tooth replica to check its

fit accuracy using a dental explorer (EXD5 explorer, Hu-Friedy, Chicago, IL, USA).

The intaglio surfaces of the all-ceramic crowns were grit blasted for 5 seconds with 50

μm Al2O3 particles using a microetcher under 2 bar pressure at the nozzle. A

bonding/silane coupling agent containing adhesive phosphate monomer (Clearfil SE

bond/Porcelain Bond Activator, Kuraray America Inc., New York, NY, USA) was

applied to the intaglio surfaces of zirconia-based crowns according to the manufacturer’s

recommendations.32

The intaglio surfaces of metal-ceramic crowns were then treated

using an alloy primer (Kuraray America Inc.).

The surfaces of molar replicas were etched using 40% phosphoric acid (Kuraray

America Inc.) for 1 minute. The etched surfaces were cleaned using water spray and

dried using oil-free compressed air.

All crowns were cemented with dual-cured phosphate-modified resin cement

(Panavia F 2, Kuraray America Inc.). ED primer (Kuraray America Inc.) was mixed and

applied evenly to the etched tooth replica surfaces. After 30 seconds, the primer was

dried with compressed air. Panavia F 2 pastes A and B were mixed for 20 seconds and

then applied to the internal surface of the crown. Initially, the crowns were seated on

their respective replicas using finger pressure, the excess cement was removed, and an air

inhibiting gel (Oxyguard, Kuraray America Inc.) was applied to the crown margin for 3

minutes to minimize the formation of the oxygen-inhibited layer at the margin. The

crowns were then placed under static pressure of 2.2 Kg for 5 minutes.33

After removing

the static load, each crown was light cured at the margin (20 seconds for each surface)

using a light curing unit (Optilux 501, Kerr Demetron, Danbury, CT, USA). One hour

after cementation, the crowns were stored in 37˚ C distilled water for 1 week.

68

Mechanical cyclic fatigue and fracture test

All cemented crowns with their respective dies were mounted in resin material (SR

Ivolen, Ivoclar Vivadent) with dimensions suitable for attachment to a loading jig. All

crowns were subjected to mechanical cyclic loading at loads ranging between 50 and 600

N for 500,000 cycles at a frequency of 20 Hz in distilled water at room temperature to

mimic the hydrolytic effect of saliva on the ceramic (static fatigue) using a universal

testing machine (Instron 8501, Instron, Canton, MA, USA). A 5-mm diameter tungsten

carbide ball was used to apply the forces along the long axis of the crown through the

central fossa. The Instron machine was adjusted to stop if the deformation increased

more than 0.15 mm. After cyclic loading, the crowns were examined using magnifying

loupes (X 2.5) for any cracks or fractures. All crowns that passed the cyclic loading

without any evidence of cracks and/or fracture were further subjected to compressive

loading in an Instron machine (Instron 8501) using a cross-head speed of 0.5 mm/min.

Fracture was defined as the occurrence of a sharp drop in the load and confirmed with

acoustic events. At the first sign of fracture, the test was terminated.

The load-displacement data from the Instron machine were imported to a software

program (Microsoft Office Excel, Redmond, WA), and load-displacement curves were

generated. Each curve was carefully analyzed to determine the failure point (the first

sharp drop in the curve) and the corresponding load (N) was recorded as the failure load.

All fractured specimens were examined under magnification (X 2.5) to determine the

mode of fracture (if it involved the veneering porcelain layer alone or both the veneering

porcelain and core layers). In addition, the extent of fracture was classified based on the

extent of the occlusal involvement (involving 50% or less, or more than 50% of the

occlusal surface) and the involvement of any of the axial surfaces (yes or no).

Four zirconia-based crowns were selected (1 from each group) for scanning electron

microscopy (SEM) examination. The crowns were sputter-coated with 7 nm of platinum

in a Polaron E5100 coating unit (Polaron Equipment, Ltd., Bedford, UK) and examined

using SEM (Hitachi S-2500, Hitachi, Mito City, Japan).

69

Statistical Analysis

Two-way analysis of variances (ANOVA) was used to evaluate the effect of core

thickness and core design on fracture strength. One-way ANOVA, followed by Tukey's

test, was used to compare the fracture loads of all groups (including the control). Chi-

square or Fisher’s exact test was used to evaluate the effect of core thickness and design

on the fracture mode and extent.

All statistical analyses were conducted using the SPSS software for Mac (version 20,

SPSS Inc., IBM, Somers, New York, USA). All statistical analyses were two-tailed at a

significance level of (0.05).

70

Results

Fatigue Resistance and Fracture Strength

All zirconia-based crowns passed the cyclic loading without any signs of cracks or

fractures, while two of the metal-ceramic crowns showed signs of cracks or fractures.

Two-way ANOVA revealed significant main effects of the core thickness (P <0.05), but

not for the core design and the interaction term (Table 4.3). The fracture strength of

crowns with 1 mm core thickness (mean [SD]=3061 [816] N) was significantly higher

than the crowns with 0.6 mm core thickness (mean [SD]=2543 [761] N).

One-way ANOVA revealed a statistically significant difference between the groups

(including the metal-ceramic crowns) (P=0.003). Tukey’s post-hoc comparisons

indicated that the fracture strength of both zirconia-based crowns with 1mm anatomical

cores (group 4) and metal-ceramic crowns (control) were statistically higher than the

zirconia-based crowns with 0.6 mm non-anatomical cores (group 1). The fracture

strength data of all groups are summarized in Table 4.4.

Mode and Extent of Fracture

Distributions of the different fracture modes among the groups are shown in Table 4.5.

Except for two, all the zirconia-based crowns (95%) showed a pattern of fracture that

involved the veneering porcelain layer only. Fisher’s exact test revealed no statistically

significant effect of the core thickness and design on the fracture mode of the zirconia-

based crowns (P>0.05). The metal-ceramic crowns revealed similar fracture mode with

all the crowns showing fracture involving the veneering layer only.

Half of the zirconia-based crowns showed fractures involving 50% or less of the

occlusal surface with chi-square test showing no significant difference among the

different core designs and thicknesses. In contrast, about two-thirds of the zirconia-based

crowns showed fracture that involved at least one of the axial surfaces with chi-square or

Fisher’s exact tests showing a statistically significant difference among the different core

designs (P<0.001), but not among the different core thicknesses. Higher percentage of

fractures that involved at least one of the axial surfaces was observed among the

71

zirconia-based crowns with non-anatomical core in comparison to the anatomical ones.

Distribution of the fracture extent data among different independent variables is shown in

Table 4.6. Some examples of the different fractures extent are presented in Figure 4.3.

Half of the metal-ceramic crowns showed fractures that involved more than 50% of

the occlusal surface and only 25% showed axial involvement.

Microscopic examination

The core material showed a homogenous structure compared to the veneering

porcelain where porosities were observed. SEM examination showed that the fracture

originated from the occlusal surface at the loading area and propagated within the

veneering porcelain. At the core-veneer interface, the crack propagated parallel to the

interface (Figure 4.4). It is not clear if the failure at the interface was cohesive within the

veneering porcelain or adhesive at the interface. Some of the SEM pictures revealed

remnants of the veneering porcelain on the core (Figure 4.5). In two of the zirconia-

based crowns, the crack propagated into the core in addition to the veneering porcelain

(Figure 4.6).

Visual examination of the metal-ceramic crowns revealed that all the crowns showed

a combined mode of fracture (cohesively within the veneering porcelain and adhesively

at the metal-ceramic interface (Figure 4.7).

72

Discussion

Despite the limitations of in-vitro studies, they allow the evaluation of individual

factors on the performance of the dental restorations under standardized conditions. In

this study, the recommendations for a clinically-relevant in-vitro load-to-failure test for

all-ceramic restorations described by Kelly were followed,34

including preparing the

teeth according to the clinical guidelines, crown fabrication according to the clinical

standards and with clinically-relevant dimensions, and using a reliable, commonly used

luting cement. The die material used had elastic modulus similar to that of dentin to

avoid variability associated with natural teeth. An aqueous environment was maintained

during cyclic-loading to resemble the effect of water on promoting the crack growth

(static fatigue).2

Clinically, all-ceramic restorations commonly fail through slow crack growth

resulting from fatigue caused by masticatory stresses. The crowns in this study were

fatigued using a range of forces that are clinically-relevant in the molar region. The

crowns were loaded for 500,000 cycles, which represents 10 years of normal function in

the oral cavity,35

or as suggested by Kelly, about half a year of continuous bruxism.34

Our study is unique in terms of examining both the effect of varying core design and

thickness on the in-vitro performance of the zirconia-based crowns. Previous studies

focused on only one of these two factors.25-29, 36

In contrast to the previous studies, no

significant effect of the core design on the fracture strength and fatigue resistance of the

zirconia-based crowns was found in the present study. This dichotomy between the

results of this current study and that of the previous studies can be attributed to different

methodology (including different core designs, materials, core to veneer thickness ratios

and loading protocols). In the present study, the core design was modified occlusally to

provide even support for the veneering layer. Other design modifications were evaluated

in the literature. Recent laboratory work has shown that a modified core design

comprising a high lingual collar connected to high strut at the proximal areas improved

the in-vitro performance of zirconia-based crowns.37-39

Future in-vitro and in-vivo studies

that assess the effect of combining different design modifications on the performance of

the zirconia-based crowns are warranted.

73

In contrast to the core design, core thickness influenced the fracture strength of the

zirconia-based crowns. Results of the present study are in agreement with those of other

studies that favored high core to veneer thickness ratio to improve the fracture strength.7,

30, 36

In the present study, almost all the zirconia-based crowns, except two, showed

fracture mode that involved the veneering layer only without core involvement, which is

expected due to the lower fracture strength and fracture toughness of the veneering

porcelain in comparison to the core. This finding is in harmony with previous clinical

studies, which reported that the cohesive failure within the veneering porcelain is the

chief failure mode for the zirconia-based crowns with the framework fracture seldom

reported.10, 40, 41

In agreement with previous in-vitro studies, core design and thickness

did not affect the fracture mode in this study.25-29, 36

Guazzato et al.42

studied the fracture mode of bilayer porcelain/zirconia samples and

found that fracture initiation occurred through the development of a Hertzian cone crack

at low force. When the force increased, this crack propagated and approached the core

with an acute angle. As the crack reached the core-veneer interface, it was hindered, and

extended laterally parallel to the interface. This lateral extension resulted in delamination

of the veneering porcelain and was followed by crushing of the contact area in 80 % of

the cases. Comparable findings were observed from the fractographic analysis of the

crowns in this study (Figures 4.4 and 4.5) and from examining failed zirconia-based

crowns clinically in a previous report 43

, with the fracture starting at the contact point and

propagating parallel to the interface. As stated previously, visual examination did not

allow the recognition of the fracture nature (cohesive or adhesive) because of the

translucency of the zirliner layer. Some SEM pictures in this current study showed

remnants of veneering material on the core layer. Similar findings were observed in a

previous study.44

The difference of the fracture extent (axially) between the different core designs is in

agreement with previous studies that reported smaller chipping sizes with modified core

designs.26, 44

As stated previously, the crack started at the contact area occlusally and

propagated parallel to the interface. In case of simple even thickness core, the crack is

expected to propagated without any deflection toward the axial surface. In contrast,

74

anatomical cores will deflect the cracks away from the axial surfaces; hence, decreasing

the size and extent of the fracture. It has to be highlighted that such reduced fracture size

translates to a more feasible chairside repair (re-polishing or repair with restoration)

should it occur clinically.

Metal-ceramic crowns were selected as a control in this study because they are

reputed to be the gold standard in dentistry, offering reasonable esthetics45

and long-term

performance.46

The fact that two of the metal-ceramic crowns fractured during fatigue

test might be an indication of their susceptibility to fracture under high occlusal forces.

Goodacre et al.1 reported that porcelain fracture is among the most common

complications for metal-ceramic single crowns at a mean rate of 3% (range between 2.7

and 6%). Despite being a common complication for single crowns, porcelain fracture is

not reported as a significant issue clinically, perhaps because most of the time it is not

noticed by the patient or it is too small in size to cause significant esthetic or functional

concerns.10

In contrast to metal-ceramic restorations, most of the studies that evaluated

the clinical performance of the zirconia-based restoration are short-term and focused

mainly on the fixed partial dental prostheses and not single crowns.10

Although zirconia-

based crowns were introduced as an alternative to metal-ceramic, most of the in-vitro and

in-vivo studies failed to include metal-ceramic as a control. A thorough literature search

identified only two short-term randomized controlled trials (12-18 months) comparing

zirconia-based to metal-ceramic crowns, with comparable performance clinically. 12

13

In the present study, higher mean fracture load was recorded for the metal-ceramic

crowns in comparison to group 1 (0.6 mm even thickness core) but not for the other

groups. This may be attributed to the higher strength of the veneering porcelain used in

the metal-ceramic, the difference in the internal residual stresses within the crowns as a

result of the higher thermal conductivity and lower modulus of elasticity of the metal

alloy in comparison to zirconia and/or improved core-veneer bond. Similar findings were

reported previously using the same zirconia and veneering porcelain but without any

details about the core design.47

Silva et al. 43

reported significantly higher reliability

under mouth-motion cyclic mechanical testing for the metal-ceramic crowns in

comparison to the zirconia-based ones.

75

Similar to the zirconia-based crowns, the metal-ceramic crowns revealed a fracture

mode originating from the occlusal contact area and spreading laterally, with signs of

adhesive fracture and exposure of the metal core, which was easier to identify in

comparison to the zirconia-based crowns due to absence of the opaque layer. Similar

mode of failure was observed previously in-vitro and in-vivo studies. 43, 44

As all loading in the study was vertical and other forces were ignored due to the

loading machine limitations, the clinical implication of the results of the present study

should be interpreted with caution. The use of a higher frequency (20 Hz) in the cyclic

loading test rather than a lower frequency (1 to 2 Hz) as observed in chewing cycles was

unavoidable because of budget limitation. Such high frequency may lead to more heat

generation compared to 1 to 2 Hz, and may not give a time for stress relaxation. Kelly et

al. used 20 Hz frequency for cyclic loading of leucite-reinforced all-ceramic crowns

using a staircase approach between 100 to 600 N, with a 100 N step size for 106 cycles in

water, and was able to measure fracture loads reasonably well for clinical relevance.48

Fracture loads were recorded when there was a sharp drop in the load-displacement

curve accompanied by acoustic sound. However, the load at complete fracture may not

necessarily be the load that the material will withstand before crack initiation. Generally,

crack initiation occurs at a lower force compared to the force required for complete

fracture. Once a crack is initiated, crack propagation starts, leading to complete fracture.

This study did not record the “pop-in” crack (crack initiation) of the specimens because

the crack initiation was not accompanied by any drop in the load, since the loading ball

was entirely supported by the intact specimen surface.

The use of a tungsten carbide loading ball increased the contact pressure in the

specimen compared to the clinical situation as the contact pressure is influenced by the

ratio of the elastic modulus of the porcelain to that of the loading ball, and by the radius

of the loading ball.49

Alternative loading methods could have included the use of a ball

with a modulus of elasticity lower than that of the tungsten carbide one, a tin sheet

between the load applicator and crown as stress breaker, or a stainless steel loading

piston with its end machined to a curvature equivalent to 40-50 mm diameter to

reproduce clinically-relevant contact pressure. However, this increase in the loading

stresses will not affect the comparison between the different test groups since all

76

specimens were loaded using similar contact pressure. Fracture load data reported in this

study should be considered as relative, not absolute values, and extrapolation of these in-

vitro strength data to clinical performance must be considered cautiously and within the

limitations of the study.2

77

Conclusions

This study highlighted the importance of core thickness and design as factors that

can influence the in-vitro performance of the zirconia-based crowns. Both null

hypotheses were partially rejected. The fracture strength and fracture extent were

significantly affected by core thickness and core design, respectively. In addition, metal-

ceramic crowns showed significantly higher mean fracture load in comparison to

zirconia-based crowns with 0.6 mm even-thickness copings. Long term clinical trials are

needed to confirm the findings of this study.

78

Acknowledgements:

This study is part of a PhD thesis submitted to the Faculty of Dentistry, University of

Toronto.

This work was supported by grants from the Faculty of Dentistry Research

Committee, University of Toronto; and Education and Research Foundation of

Prosthodontics. Materials donations were received from Ivoclar-Vivadent and Kuraray;

and technical support provided by Shaw Dental Labs (Toronto) to whom the authors are

grateful.

79

(Table 4.1) Composition and some properties of materials used for core and veneering porcelain

(as reported by manufacturer).


Technique

CTE

(10-6.K-1)

Flexure strength

(MPa)

IPS e.max Zirpress

(Lot: H22305)

Ivoclar Vivadent,


Fluorapatite glass-

ceramic Pressing

9.75 0.25

(100 - 400 C)

110 10

IPS e.max Zirliner

(Lot: L29047)

Ivoclar Vivadent,


Fluorapatite glass-

ceramic

Manual

application - -

IPS e.max ZirCad

(Lot: L10846)

Ivoclar Vivadent,


Yttrium-stabilized


10.75 0.25

(100 -400 C)

900 50

Authentic Alloy HE

(Batch: 12372)

Ceramay, Stuttgart,

Germany

33% Co, 30% Cr,

29% Fe and 5% Mo Casting

15.9

(20 -500 C)

-

Authentic Ingot

(Batch: 12372)

Ceramay, Stuttgart,

Germany Leucite-reinforced Pressing

15.8

(20 -500 C)

148

80

(Table 4.2) Firing programs for materials used in the study

Step

Pre drying

temperature

(ºC)

Hold time

(min.)

Temperature

increasing rate

(ºC/min)

Final firing

temperature

(ºC)

Holding time

(min)

Vacuum 1

(ºC)

Vacuum 2

(ºC)


IPS e.max ZirPress

(200g ring) 700 - 60 910 15 500 910

Authentic Ingot

(200g ring) 700 - 60 940 20 730 940

81

(Table 4.3) Two-way ANOVA

(dependent variable: fracture load; independent variables: core thickness and design)

Source Type III Sum

of Squares df Mean Square F Sig.

Corrected Model 4900197.3 3 1633399.1 2.741 0.057

Intercept 314059368.1 1 314059368.1 527.014 <0.001

Core Thickness 2691534.4 1 2691534.4 4.517 0.04

Core Design 2180890 1 2180890 3.66 0.064

Core Thickness * Core Design 27772.9 1 27772.9 0.047 0.83

Error 21453192.6 36 595922.017

Total 340412758 40


R2 = 0.186 (Adjusted R2 = 0.118)

82

(Table 4.4) Fracture load (N) data for all groups.

Group *

Core

Thickness

(mm)

Veneer

Thickness

(mm)

Mean (SD) Median

95 %

Confidence

Interval

Min - Max

Group 1: 0.6 mm non-anatomical cores a 0.6 0.9-1.4 2336 (792) 2232 1769 - 2902 1046 - 3783

Group 2: 0.6 mm anatomical cores a, b 0.6-1.1 0.9 2750 (708) 2767 2243 - 3256 1770 – 4000

Group 3: 1 mm non-anatomical cores a, b 1 0.5-1 2802 (584) 2842 2384 - 3219 1599 - 3574

Group 4: 1 mm anatomical cores b 1-1.5 0.5 3321 (956) 3507 2637 - 4005 1757 – 4500

Control: Metal-ceramic b 0.6-1.1 0.9 3722 (612) 3859 3210 - 4233 2884 - 4500

* Identical letters indicate no statistical difference based on Tukey’s test.

83

(Table 4.5) Effect of different independent variables on fracture mode

Core Thickness Total

0.6 mm 1 mm

Failure Mode (N (%)) Failure Mode (N (%)) Failure Mode (N (%))

Veneer

Only

Both

Veneer and

Core

Veneer

Only

Both

Veneer and

Core

Veneer

Only

Both

Veneer and

Core

Core Design

Non-anatomical 10 (100 %) 0 9 (90 %) 1 (10 %) 19 (95 %) 1 (5 %)

P>0.05*

Anatomical 10 (100 %) 0 9 (90 %) 1 (10 %) 19 (95 %) 1 (5 %)

Total 20 (100 %) 0 18 (90 %) 2 (10%)

P>0.05*

* using Fisher’s exact test.

84

(Table 4.6) Effect of different independent variables on fracture extent

Core

Thickness Core Design

Occlusal involvement

Significance

Level

Axial Fracture

Significance

Level 50% or

less

More than

50% No Yes

0.6 mm

Non-anatomical 3 (30 %) 7 (70 %)

P>0.05*

1 (10 %) 9 (90 %)

P=0.02*

Anatomical 7 (70 %) 3 (30 %) 7 (70 %) 3 (30 %)

1 mm

Non-anatomical 5 (50 %) 5 (50 %)

P>0.05*

0 10 (100 %)

P=0.003**

Anatomical 5 (50 %) 5 (50 %) 7 (70 %) 3 (30 %)

* using chi-square test

** using Fisher’s exact test

85

(MD: Mesio-distal and BL: bucco-lingual).

(Figure 4.1) Preparation dimensions of the master die (buccal view)

86

(Figure 4.2) Crown Fabrication

a) Different core designs

(Group 1: Non-anatomical copings with even thickness of 0.6 mm occlusally and circumferentially, Group 2:

Anatomically shaped copings with a minimum thickness of 0.6 mm occlusally and circumferentially, Group 3:

Non-anatomical copings with even thickness of 1 mm occlusally and 0.6 mm circumferentially and Group 4:

Anatomically shaped copings with a minimum occlusal thickness of 1 mm and 0.6 mm circumferential thickness).

b) Veneering layer CAD-waxx replica sealed on the zirconia coping.

c) Zirconia-based crowns after polishing.

d) Coping CAD-waxx replica for the metal-ceramic crowns

e) Metal coping after casting.

f) 1: Metal coping after the opaque layer application.

2: Veneering layer CAD-waxx replica for the metal-ceramic crowns.

3: Veneering layer CAD-waxx replica sealed on the metal coping.

g) Metal-ceramic crowns after polishing.

87

(Figure 4.3) Examples of different fracture extent

≤50% of the occlusal

surface without axial

involvement

≤50% of the occlusal

surface with axial

involvement

>50% of the occlusal

surface without axial

involvement

>50% of the occlusal

surface with axial

involvement

88

(Figure 4.4) SEM photomontage showing a fractured crown from group 2 (fracture involved the veneering layer only).

The fractographic examination revealed that the fracture started occlusally at the loading area with a cone crack and propagated parallel to the

interface toward the axial area. Crushing of the veneer ceramic can be observed at the loading area. Wake hackle markings were observed on pores as

an outcome of fracture passage around the pores. The markings indicate the direction of the crack propagation. (Zr: zirconia, V: veneer, and dcp:

direction of crack propagation)

89

(Zr: zirconia and V: veneer)

(Figure 4.5) SEM picture for a crown from group 4 showing remnant of the veneering porcelain attached to

the zirconia core.

90

(Figure 4.6) SEM photomontage showing a fractured crown from group 4 (fracture involved the veneering and core layers).

The fractographic examination revealed that the fracture started occlusally at the loading area with a cone crack and propagated into the core.

Crushing of the veneer ceramic can be observed at the loading area. Arrest lines markings indicate the direction of crack propagation . (Zr: zirconia, V:

veneer, and dcp: direction of crack propagation).

91

(Figure 4.7) Fracture mode observed in metal-ceramic crowns.

92

References




2004;48(2):viii, 513-viii, 30.

3. Anusavice KJ, Phillips RW. Phillips' science of dental materials. 10th

ed.

Philadelphia: Saunders; 1996. p. 615.


2008;24(3):299-307.



J Dent 2004;32(2):91-9.






2005;94(2):125-31.



2012;107(3):170-7.








2009;37(9):731-6.









93







Dent Mater 2009;25(7):857-62.



Prosthodont 2007;20(4):383-8.




2006;96(4):237-44.



Mater 2005;21(10):984-91.




22. Marchack BW, Futatsuki Y, Marchack CB, White SN. Customization of milled

zirconia copings for all-ceramic crowns: a clinical report. J Prosthet Dent

2008;99(3):169-73.

23. Zarone F, Russo S, Sorrentino R. From porcelain-fused-to-metal to zirconia:

clinical and experimental considerations. Dent Mater 2011;27(1):83-96.

24. Wang H, Pallav P, Isgro G, Feilzer AJ. Fracture toughness comparison of three

test methods with four dental porcelains. Dent Mater 2007;23(7):905-10.

25. Sundh A, Sjogren G. A comparison of fracture strength of yttrium-oxide-

partially-stabilized zirconia ceramic crowns with varying core thickness, shapes

and veneer ceramics. J.Oral Rehabil. 2004;31(7):682-88.

26. Guess PC, Bonfante EA, Silva NR, Coelho PG, Thompson VP. Effect of core

design and veneering technique on damage and reliability of Y-TZP-supported

crowns. Dent Mater 2013;29(3):307-16.

27. Kokubo Y, Tsumita M, Kano T, Fukushima S. The influence of zirconia coping

designs on the fracture load of all-ceramic molar crowns. Dent Mater J

2011;30(3):281-5.

28. Larsson C, El Madhoun S, Wennerberg A, Vult von Steyern P. Fracture strength

of yttria-stabilized tetragonal zirconia polycrystals crowns with different design:

an in vitro study. Clin Oral Implants Res 2012;23(7):820-6.

94



Dent 2009;37(12):978-83.








33. Pallis K, Griggs JA, Woody RD, Guillen GE, Miller AW. Fracture resistance of

three all-ceramic restorative systems for posterior applications. J Prosthet Dent

2004;91(6):561-9.



35. Wiskott HW, Nicholls JI, Belser UC. Stress fatigue: basic principles and

prosthodontic implications. Int J Prosthodont 1995;8(2):105-16.

36. Alhasanyah A, Vaidyanathan TK, Flinton RJ. Effect of Core Thickness

Differences on Post-Fatigue Indentation Fracture Resistance of Veneered

Zirconia Crowns. J Prosthodont 2013.

37. Bonfante EA, da Silva NR, Coelho PG, Bayardo-Gonzalez DE, Thompson VP,

Bonfante G. Effect of framework design on crown failure. Eur J Oral Sci

2009;117(2):194-9.

38. Bonfante EA, Rafferty B, Zavanelli RA, Silva NR, Rekow ED, Thompson VP, et

al. Thermal/mechanical simulation and laboratory fatigue testing of an alternative

yttria tetragonal zirconia polycrystal core-veneer all-ceramic layered crown

design. Eur J Oral Sci 2010;118(2):202-9.

39. Silva NR, Bonfante EA, Rafferty BT, Zavanelli RA, Rekow ED, Thompson VP,

et al. Modified Y-TZP core design improves all-ceramic crown reliability. J Dent

Res 2011;90(1):104-8.

40. Al-Amleh B, Lyons K, Swain M. Clinical trials in zirconia: a systematic review. J

Oral Rehabil 2010;37(8):641-52.

41. Guess PC, Schultheis S, Bonfante EA, Coelho PG, Ferencz JL, Silva NR. All-

ceramic systems: laboratory and clinical performance. Dent Clin North Am

2011;55(2):333-52, ix.

42. Guazzato M, Proos K, Quach L, Swain MV. Strength, reliability and mode of

fracture of bilayered porcelain/zirconia (Y-TZP) dental ceramics. Biomaterials

2004;25(20):5045-52.

43. Silva NR, Bonfante EA, Zavanelli RA, Thompson VP, Ferencz JL, Coelho PG.

Reliability of metalloceramic and zirconia-based ceramic crowns. J Dent Res

2010;89(10):1051-6.

95

44. Lorenzoni FC, Martins LM, Silva NR, Coelho PG, Guess PC, Bonfante EA, et al.

Fatigue life and failure modes of crowns systems with a modified framework

design. J Dent 2010;38(8):626-34.

45. Donovan TE. Porcelain-fused-to-metal (PFM) alternatives. J Esthet Restor Dent

2009;21(1):4-6.

46. Napankangas R, Raustia A. Twenty-year follow-up of metal-ceramic single

crowns: a retrospective study. Int J Prosthodont 2008;21(4):307-11.

47. Augstin-Panadero R, Fons-Font A, Roman-Rodriguez JL, Granell-Ruiz M, del

Rio-Highsmith J, Sola-Ruiz MF. Zirconia versus metal: a preliminary

comparative analysis of ceramic veneer behavior. Int J Prosthodont

2012;25(3):294-300.

48. Kelly JR, Hunter BD, Brenyo MR, Peterson IM. Simulating clinical failure

during in vitro testing of all-ceramic crowns. J Dent Res 1998;77:778.



96


97

Effect of core design and veneering material on fracture strength and fatigue

resistance of zirconia molar crowns

Authors:




Saudi Arabia






Toronto, ON, Canada



Toronto, ON, Canada


Associate Professor and Head, Department of Biomaterials Science, Schulich School of Medicine

& Dentistry, The University of Western Ontario, London, ON, Canada






Phone: (416) 979-4934 X 4572

Fax: (416) 979-4936


This work was previously presented in part at the 91st General Session & Exhibition of the IADR,

Seattle, WA, USA in 2013

98

Abstract

Statement of Problem: Chipping of zirconia-based crowns is a common problem.

Purpose: To determine the effect of core design and veneering material on fatigue

resistance and fracture strength of zirconia-based crowns veneered with CAD/CAM

manufactured ceramic. Material and Methods: Forty zirconia cores were fabricated

using uniform thickness or anatomically shaped designs. Cores were veneered using

CAD/CAM manufactured lithium disilicate (IPS e.max CAD) (EMAX) or feldspathic

(Vita Mark II) (VMII) ceramic veneers. Four groups were created by combining different

veneering materials and core designs (n=10). The veneers were connected to the copings

using resin cement or fusion glass-ceramic for the VMII and EMAX crowns, respectively.

All crowns were cemented onto their respective epoxy dies using a resin cement and then

subjected to compressive-mechanical fatigue in a universal testing machine. Specimens

that survived the fatigue test were loaded to fracture at crosshead speed of 1mm/min and

fracture load and mode were recorded. Data were statistically-analyzed using two-way

Analysis of Variance (ANOVA) , and chi-square and Fisher’s exact tests at α=0.05.

Results: All crowns survived the fatigue test without developing cracks or fractures.

Two-way ANOVA revealed a statistically significant effect of the veneering material

(P<0.001) on the fracture load, but not for the core design or the interaction term. The

VMII-veneered crowns showed a lower mean fracture load in comparison to the EMAX -

veneered ones. All VMII-veneered crowns showed fractures involving the veneer only,

while 70% of the EMAX crowns had fractures involving both the core and veneer.

Conclusion: The veneering material, but not the core design, had a significant effect on

fracture load and mode of zirconia-based crowns.

Keywords: Fracture strength; fracture mode: zirconium oxide ceramics; core design;

CAD/CAM

99

Introduction



alternatives to metal-ceramic crowns such as all-porcelain ones. However, the clinical

performance of all-ceramic restorations stood short of the ideal due to their inherent



In a systematic review on clinical complications in fixed prosthodontics,1 all-ceramic

crowns showed an 8% incidence of complications, with crown fractures being the most

common. Molars showed a higher fracture rate (21%) than premolars and anterior teeth

(7% and 3%, respectively). Kelly reported similar results with higher fracture rates in

posterior crowns compared to anterior crowns.2 This higher fracture rate in the posterior

region can be explained by the higher biting forces in the posterior region (400 to 890 N

in the molar region and 222 to 445 N in the premolar area) in comparison to the anterior

region (133 to 334 N in the cuspid area and 89 to 111 N in the incisor region).3

In the past decade, dental manufacturers have focused their efforts on developing

new all-porcelain products with enhanced mechanical properties to improve their clinical

performance in the posterior region through enhanced resistance to chipping and fracture.

Yttrium-tetragonal zirconia poly-crystals (Y-TZP) is one of these products that has

received special attention due to its high flexural strength and fracture toughness.4 The

tetragonal crystals in these zirconium oxide ceramics are meta-stable and can be

transformed into larger monoclinic crystals with the application of stress from cracks or

flaws.5 This phenomenon is beneficial in hindering crack growth and increasing fracture

toughness; hence, it is referred to as “transformation toughening”.6 However, Y-TZP-

based products cannot provide optimal esthetics because they possess high opacity and

are difficult to stain. Therefore, Y-TZP must be used as a core material and clad with

veneering porcelain to enhance esthetics. 7

At the early stages in the history of zirconia-based crowns, veneering porcelain was

layered manually on the core material and sintered similar to metal-ceramic crowns. The

improved mechanical properties of Y-TZP -based restorations were not readily reflected

100

in their clinical performance with several short-term clinical studies reporting a high

incidence of veneering porcelain chipping with rare framework fractures.8-14

Although

minor veneering porcelain chipping, in most of the cases, does not necessitate

replacement of the restoration, it may create a rough surface or sharp edge causing

patient discomfort or if involving the esthetic zone, it may compromise the esthetics.

This high chipping rate, in comparison to that of metal-ceramic restorations may be

caused by: occlusal overloading, residual stresses as a result of the mismatch in the

coefficient of thermal expansion (CTE) between core and veneering porcelain, 15, 16

residual “tempering” stresses due to the thermal gradient between the veneering

porcelain and the zirconia core,16

interfacial stresses as a result of crystal transformation

at the interface,17


inherent strength deficiency of the

veneering porcelain 18, 19


Several potential

solutions were suggested in clinical and in-vitro studies to improve the clinical

performance of zirconia-based restorations. These were mainly directed towards

improving the strength of the veneering layer and the veneer-to-core bond. Two of these

suggestions were directed towards enhancing the core design to provide adequate and

even support to the veneering porcelain,22

and improving the strength of the veneering

layer. 23, 24

In general, the ability of ceramic material to withstand occlusal forces is

compromised by the presence of two types of inherent flaws25

: fabrication defects

(internal voids, porosities or microstructural features which arise during processing) and

surface cracks (defects on the surface as a result of machining and grinding process).26

Failure begins with microscopic damage that results from the interaction of preexisting

defects with applied loads.25

Failure can also occur because of impact forces or

subcritical crack growth,27, 28

which is enhanced in an aqueous environment.29

Thus, it is

expected that reducing the flaws and voids within the veneering material may result in

better mechanical properties.

In comparison to the manual layering, heat pressed veneering materials with thermal

properties compatible with the zirconium oxide cores improved the strength and the

interfacial bond between the veneering porcelain and zirconia.23

This improvement was

reflected in the clinical performance of zirconia-based restorations with short-term

101

clinical studies showing significantly less veneer chipping for pressed veneering

porcelain than for manually applied veneering porcelain.10

Unfortunately, the

improvement in the performance of the veneering layer is limited by the flexural strength

of the veneering layer, which is generally less than 125 MPa.30, 31

A potential alternative route is to fabricate the veneering and core layers

independently by means of CAD/CAM technology using industrially fabricated ceramic

blocks and then join them using fusion ceramic or adhesive cement.32, 33

This approach

simplifies the fabrication process, saves time by eliminating time-consuming hand-

layering, drying, and fusing processes, and allows the use of high strength ceramic that is

free from processing flaws as a veneering layer, which is expected to improve the

strength of the zirconia-based restorations. IPS e.max CAD-on (Ivoclar Vivadent,

Schaan, Liechtenstein) and Vita Rapid Layer (VitaZahnfabrik, BadSackingen, Germany)

use lithium disilicate (IPS e.max CAD, Ivoclar Vivadent) and feldspathic (Vita Mark II,

VitaZahnfabrik) blocks, respectively, to produce the veneering layer for all-ceramic

zirconia-based restorations using CEREC system (Sirona Dental Systems GmbH,

Bensheim, Germany). In both techniques, the core and the veneering layers are designed

and fabricated using the CAD-CAM technology and connected using a resin cement (for

the rapid layering technique) and a fusion glass-ceramic (for the CAD-on technique).

Several in-vitro studies evaluated the effect of core design modification as a factor

that might influence the performance of zirconia-based restorations. 34-38

These studies

used different methodologies and provided conflicting results. Clinical studies on

zirconia fixed partial dentures with anatomic framework design showed promising results,

but these studies were short-term case series without controls.39, 40

Thus, the aim of this study was to evaluate the influence of different core designs

and veneering materials on the fatigue resistance, fracture strength and fracture mode of

CAD-CAM zirconia-based crowns. The objective was to test the following null

hypothesis: There is no significant difference in fatigue resistance, fracture strength and

fracture mode of zirconia-based crowns fabricated using two different zirconia core

designs (uniform thickness and anatomically shaped) and two veneering materials (IPS

e.max CAD and Vita Mark II).

102

Material and Methods

Preparation

An ivorine mandibular molar replica was prepared to receive an all-ceramic crown

following dimensions recommended for the CEREC system (Sirona Dental Systems

GmbH, Bensheim, Germany) (Figure 5.1). The tooth had an occlusal reduction of 1.5

mm and minimum axial reduction of 1 mm with a 6 to 8° total angle of convergence (3 to

4° on each side). The gingival margin was prepared with a circumferential shoulder at

least 1 mm wide. Lingual and facial surfaces were prepared in two planes, and all line

angles were smoothed to reduce stress concentration. The prepared ivorine molar was

used as a master die to fabricate 40 replicas using a highly filled epoxy resin (Viade

Products, Inc., Camarillo, CA) with a modulus of elasticity similar to that of human

dentin (12.9 GPa). This material responded to 34% phosphoric acid etching by forming

surface microroughness suitable for bonding. 41

Coping design and fabrication

The prepared molar replica was inserted in a dentiform with mesial and distal

adjacent ivorine teeth in place. The surfaces were then covered with a thin layer of

optical reflective medium (CEREC Optispray, Sirona Dental Systems GmbH). A scan of

the prepared tooth with its adjacent teeth was then captured using the inEos Blue scanner

(Sirona Dental Systems GmbH). A reference master crown, which had a minimum

thickness of 1.5 mm at the central fossa and 2mm at the cusp area, was designed and

milled using the biogeneric mode of the inLab software (Version 3.85, Sirona Dental

Systems GmbH), and Vita Mark II blocks (VitaZahnfabrik, BadSackingen, Germany)

using the inLab MCXL milling unit (Sirona Dental Systems GmbH).

Two different coping designs (Figures 5.2a, 5.2b and 5.2c) were used with the inLab

software (Version 3.85) as follows:

Non-anatomical copings with a uniform thickness of 0.6 mm occlusally and


103

Anatomically-shaped copings with a minimum thickness of 0.6 mm

occlusally and circumferential (group 2).

The non-anatomical copings were designed using the multilayer mode. In this mode,

a fully anatomical crown was designed first using the master crown as a reference for the

occlusal anatomy. Then, the inLab software automatically split the design file into two

separate files for the core and veneer layers. The software produced a non-anatomical

copings design with a predefined thickness of 0.6 mm and did not allow any modification

of the design. Due to this limitation, a different designing mode (reduction mode) was

used to design the anatomical copings using the master crown as a reference for the

occlusal anatomy. The thickness of the master crown was reduced by 0.9 mm to produce

anatomical copings with minimum thickness of 0.6 mm. All copings were designed with

a 0.5 mm circumferential collar and milled from Y-TZP (IPS e.max ZirCAD blocks, size

C15, Ivoclar Vivadent) using the inLab MCXL milling unit (Sirona Dental Systems

GmbH). All copings were dried and sintered according to the manufacturer’s instructions.

The sintering procedure was conducted in a sintering furnace (Programat S1, Ivoclar

Vivadent) at a temperature of 1530◦C. The duration of the program, including the cooling

phase, was approximately 90 minutes.

Veneering layer fabrication.

The copings were veneered using CAD/CAM manufactured lithium disilicate (IPS

e.max CAD) (EMAX) or feldspathic (Vita Mark II) (VMII) ceramic veneers. The

multilayer mode was used to design the veneering layer for the non-anatomical copings

with a minimal of 0.9 mm at the central fossa. Different designing mode was used to

design the veneering layer for the anatomical copings. First, one of the anatomical

copings was inserted on the prepared molar replica. Then, the surfaces of the copings

were covered with a thin layer of optical reflective medium (CEREC Optispray) and

scanned using the inEos Blue scanner. The inLab software was used to design the

veneering layer with a uniform thickness of 0.9 mm using the master crown as reference

for the occlusal anatomy. The veneering layers were milled using EMAX and the VMII

blocks in the inLab MCXL milling unit. After milling, the copings from each design

104

were randomly assigned to two veneering materials (n=10) and each veneering layer was

fitted on its respective coping.

Crown fabrication.

All the crowns in this study were fabricated by the primary author to eliminate

variability.

The VMII veneering layers were glazed using Vita Glaze-LT (Vita Zahnfabrik)

according to the manufacturer’s instructions using a porcelain furnace (Vita vacumat

4000 premium T, Vita Zahnfabrik) (5.2d). The intaglio surfaces were treated with 5%

hydrofluoric acid-etching gel (Ivoclar Vivadent) for 20 seconds. The etched internal

surfaces were rinsed with water spray, followed by ultrasonic cleaning in distilled water

for 60 seconds. A bonding/silane coupling agent containing adhesive phosphate

monomer (Clearfil SE bond/porcelain bond activator, Kuraray America, Inc.) was then

applied to the intaglio surfaces of the VM II veneers. The external surfaces of the

zirconia copings were grit-blasted for 5 seconds with 50 μm Al2O3 particles using a

microetcher under 2.5 bar pressure.

The VM II veneering layers were cemented onto their respective copings with dual-

polymerized phosphate-modified resin cement (Panavia F 2.0, Kuraray America, Inc.)

according to manufacturer’s instructions. The veneers were seated on their respective

copings using finger pressure, the excess cement was removed, an air-sealing gel

(Oxyguard, Kuraray America, Inc.) was applied to the junction between the coping and

the veneering layer for 3 minutes, and then each crown was light-polymerized (Optilux

501, Kerr Demetron, Danbury, CT) for 20 seconds on each surface.

The EMAX veneers (Figure 5.2e) were fused to their respective copings using a

fusion glass-ceramic material (IPS e.max CAD Crystall/Connect, Ivoclar Vivadent),

which was applied to the intaglio veneer surfaces and dispersed using a vibration device

(Ivomix, Ivoclar Vivadent). The copings were seated on the veneering layer using finger

pressure and the Ivomix device, which turns the fusion glass-ceramic mass into a

flowable material, and the excess fusion material was removed. The crowns were cleaned

with a brush and glaze was applied (IPS e.max CAD Crystall./glaze, Ivoclar Vivadent)

105

(Figure 5.2f). The fusion and crystallization firing as well as the glaze firing were

completed according to the manufacturer’s recommendations in a porcelain furnace

(Programat P500, Ivocalr Vivadent).

Details about the materials used in this study and their firing protocols are listed in

Tables 5.1 and 5.2, respectively.

Crown cementation

Before cementation, the crown thickness at the central fossa was verified to be 1.5

mm using a caliper (Buffalo Dental Manufacturing Co., Syosset, NY, USA) and the

crown was seated on its respective tooth replica to check its fit accuracy using a dental

explorer (EXD5 explorer, Hu-Friedy, Chicago, IL, USA).

The intaglio surfaces of the crowns were grit blasted for 5 seconds with 50 μm Al2O3

particles using a microetcher under 2 bar pressure at the nozzle. A bonding/silane

coupling agent containing adhesive phosphate monomer (Clearfil SE bond/Porcelain

Bond Activator, Kuraray America Inc., New York, NY, USA) was applied to the intaglio

surfaces of zirconia-based crowns according to the manufacturer’s recommendations.42

The surfaces of molar replicas were etched using 40% phosphoric acid (Kuraray

America Inc.) for 1 minute. The etched surfaces were cleaned using water spray and

dried using oil-free compressed air.

All crowns were cemented with dual-cured phosphate-modified resin cement

(Panavia F 2.0, Kuraray America Inc.) according to the manufacturer’s instructions. ED

primer (Kuraray America Inc.) was mixed and applied evenly to the etched tooth replica

surfaces. After 30 seconds, the primer was dried with compressed air. Panavia F 2.0

pastes A and B were mixed for 20 seconds and then applied to the internal surface of the

crown. Initially, the crowns were seated on their respective replicas using finger pressure,

the excess cement was removed, and the air inhibiting gel (Oxyguard, Kuraray America

Inc.) was applied to the crown margin for 3 minutes. The crowns were then placed under

static pressure of 2.2 Kg for 5 minutes.43

After removing the static load, each crown was

light cured at the margin (20 seconds for each surface) using a light curing unit (Optilux

106

501, Kerr Demetron, Danbury, CT, USA). One hour after cementation, the crowns were

stored in 37˚ C distilled water for 1 week.

Mechanical cyclic fatigue and fracture test

All cemented crowns with their respective dies were mounted in resin material (SR

Ivolen, Ivoclar Vivadent) with dimensions suitable for attachment to a loading jig.

Crowns were subjected to compressive cyclic loading with a load cycle of 50 and 600 N

for 500,000 cycles at a frequency of 20 Hz in distilled water at room temperature (Instron

8872, Instron, Canton, MA, USA). This was followed in order to mimic the hydrolytic

effect of saliva on the ceramic (static fatigue). A 5-mm diameter tungsten carbide ball

was used to apply the forces uniaxially through the central fossa. The Instron machine

was adjusted to stop if the deformation increased more than 0.15 mm. After cyclic

loading, the crowns were examined using magnifying loupes (X 2.5) for any cracks or

fractures. All crowns that passed the cyclic loading test without any evidence of cracks

and/or fracture were uniaxially-loaded in a universal testing machine (Instron 3345,

Instron) using a cross-head speed of 0.5 mm/min to fracture. Fracture was defined as the

occurrence of a sharp drop in the load and confirmed with acoustic events. At the first

sign of fracture, the test was terminated and the corresponding load (N) was recorded as

the failure load.

All fractured specimens were examined using magnifying loupes (X 2.5) to

determine the mode of fracture (if it involved the veneering porcelain layer alone or both

the veneering porcelain and core layers).

Two of the CAD crowns, representing different fracture modes, were selected for

scanning electron microscopy (SEM) examination. The crowns were sputter-coated with

7 nm of platinum in a Polaron E5100 coating unit (Polaron Equipment, Ltd., Bedford,

UK) and examined using SEM (Hitachi S-2500, Hitachi, Mito City, Japan).

107

Statistical Analyses

Two-way analysis of variances (ANOVA) was used to evaluate the effect of core

design and veneering layer material on fracture load, while Chi-square or Fisher’s exact

test was used to evaluate the effect of veneering material and core design on the fracture

mode.

All statistical analyses were conducted using the SPSS software for Mac (version 20,

SPSS Inc., IBM, Somers, New York, USA). All statistical analyses were two-tailed at a

significance level of (0.05).

108

Results

Fracture Mode

Distributions of the different fracture modes among the groups are shown in Table

5.3. All the VMII crowns showed a pattern of fracture that involved the veneering

porcelain layer only (Figure 5.3a), while only 30% of the EMAX crowns showed similar

fracture mode (Figure 5.3b) with the majority of the EMAX crowns showing a fracture

mode that involved both the veneer and core layers (Figure 5.3c). Chi-square test

revealed no statistically significant effect of the core design, but a significant effect of the

veneering material on the fracture mode.

Fatigue Resistance and Fracture Load

All crowns passed the cyclic loading test without any signs of cracks or fractures.

Two-way ANOVA revealed a significant main effect of the veneering layer (P <0.001),

but not for the core design or the interaction term (Table 5.4). Mean fracture strengths of

the EMAX crowns (mean [SD]=3545 [1116] N, 95% CI: 3022 - 4067) were significantly

higher than those of the VMII crowns (mean [SD]=2016 [434] N, 95% CI: 1813 - 2219).

The fracture strength data of all groups are summarized in Table 5.3.

Because EMAX crowns showed different fracture modes, the fracture strength data

were examined according the fracture mode (Table 5.5). The mean fracture load for the

crowns that showed fractures involving the veneering layer only was lower than the

crowns that showed fractures involving both the core and veneer layers, however, data

were overlapping between both fracture modes.

Microscopic examination

In this study, both veneering materials showed more homogenous structures with

less flaws and porosity in comparison to the manually-applied and pressed veneering

porcelain as observed in previous studies.44, 45

SEM examination showed some flaws at

the interface between the fusion ceramic and the core material (Figure 5.4). Visual and

SEM examinations showed that the fracture originated from the occlusal surface at the

109

loading area and propagated within the veneering porcelain for both veneer layer

materials (Figure 5.5). SEM photos revealed no obvious signs of crack penetration

through the interface toward the zirconia, which may indicate that the core fracture that

was observed in the majority of the EMAX crowns was due to radial cracks starting at

the cement-core interface.

110

Discussion

Despite the limitations of in-vitro studies, they allow the evaluation of individual

factors on the performance of the dental restorations under standardized conditions. In

this study, the recommendations for a clinically-relevant in-vitro load-to-failure test for

all-ceramic restorations described by Kelly were followed,46

including preparing the

teeth according to the clinical guidelines, crown fabrication according to the clinical

standards and with clinically-relevant dimensions, and using a reliable, commonly used

luting cement. The die material used had elastic modulus similar to that of dentin to

avoid variability associated with natural teeth. An aqueous environment was maintained

during cyclic-loading to resemble the effect of water on promoting the crack growth

(static fatigue).2

Clinically, all-ceramic restorations commonly fail through slow crack growth

resulting from fatigue caused by masticatory stresses. The crowns in this study were

fatigued using a range of forces that are clinically-relevant in the molar region. The

crowns were loaded for 500,000 cycles, which represents 10 years of normal function in

the oral cavity,47

or as suggested by Kelly, about half a year of continuous bruxism.46

In contrast to the previous studies, no significant effect of the core design on the

fracture load of the zirconia-based crowns was found in the present study. 34-38, 48

This

dichotomy between the results of the current study and those of the previous studies can

be attributed to different methodology (including different core designs, materials, core

to veneer thickness ratios and loading protocols). Our group observed similar finding in a

previous study (unpublished data) using similar methodology (including similar core

design, crown dimensions and loading protocol), but pressed veneering porcelain.

Crowns made using the EMAX veneer material showed higher mean fracture load in

comparison to VMII. This result was expected due to the improved mechanical

properties of the lithium disilicate in comparison to feldspathic ceramic (Table 5.1). The

different methods for joining the VMII and EMAX veneers to the cores may also have

contributed to the different fracture strength results for these groups.

In the present study, different fracture modes were observed between the different

veneering layer materials and a trend of a greater incidence of both veneer and core

111

fracture mode occurring with higher fracture loads was observed. While all the VMII

crowns showed fractures that involved the veneering layer only without core

involvement, about 70% of the EMAX crowns showed fractures involving both core and

veneer layers. The involvement of both layers in the EMAX crowns reflects a

competition between the cone crack within the veneer layer and radial crack from the

undersurface of the core layer. The increase in the strength and fracture toughness of the

EMAX veneer material over the VMII material resulted in an increase of the critical load

for the cone cracking relative to the radial cracking, which rendered the crowns more

susceptible to radial cracking of the core layer 49

. One of the advantages of adhesive

joining of the veneer and core layers is the elimination of residual thermal stresses within

the veneering layer that can be generated as a result of thermal gradients occurring during

the cooling of the veneer during fabrication, and the mismatch in thermal expansion

properties between the core and veneering porcelain.50

This advantage might have been

masked in this study as a result of the differences in the mechanical properties of the used

veneering ceramics. Future in-vitro studies that evaluate the effect of joining lithium

disilicate to zirconia core using adhesive cement or fusion ceramic on the fracture

strength and mode are warranted.

The use of adhesive cement to bond the veneer and core layers may act as a weak

interface that is subject to chemical degradation. It was shown that the adhesive cement

provides an internal barrier to crack propagation from one layer to the next, leading to

fracture modes including delamination at the interface and radial cracking at the bottom

of the flexing veneer layer.32

In the present study, cone cracks were observed within all

the VMII layers (rapid layering technique) (Figure 5.3a). It is not clear whether the

failure of the crown resulted from these cone cracks or other radial cracks at the

undersurface of the veneering layer. Hermann et al.51

evaluated the effect of cyclic

loading on the failure mode of glass-zirconia bilayered specimens and reported that cone

crack was the cause of failure in most of the cases. In contrast, Lee et al. 32

reported that

radial cracking is the cause of failure for the most of the glass-zirconia bilayered

specimens subjected to load to failure test.

Since loading in the study was vertical and other forces that are generated during

mastication were not replicated due to the loading machine limitations, the clinical

112

implication of the results of the present study should be interpreted with caution. The use

of a higher frequency (20 Hz) in the cyclic loading test rather than a lower frequency (1

to 2 Hz) as observed in chewing cycles was unavoidable because of budget limitations.

Such high frequency may lead to more heat generation compared to 1 to 2 Hz, and may

not allow sufficient time for stress relaxation. Kelly et al. used 20 Hz frequency for

cyclic loading of leucite-reinforced all-ceramic crowns using a staircase approach

between 100 to 600 N, with a 100 N step size for 106 cycles in water, and was able to

measure fracture loads reasonably well for clinical relevance.52

Fracture loads were recorded when there was a sharp drop in the load-displacement

curve accompanied by acoustic sound. However, the load at complete fracture may not

necessarily be the load that the material will withstand before crack initiation. Generally,

crack initiation occurs at a lower force compared to the force required for complete

fracture. Once a crack is initiated, crack propagation starts, leading to complete fracture.

This study did not record the “pop-in” crack (crack initiation) of the specimens because

the crack initiation was not accompanied by any drop in the load, since the loading ball

was entirely supported by the intact specimen surface.

The use of a tungsten carbide loading ball increased the contact pressure in the

specimen compared to the clinical situation as the contact pressure is influenced by the

ratio of the elastic modulus of the porcelain to that of the loading ball, and by the radius

of the loading ball.53

Alternative loading methods could have included the use of a ball

with a modulus of elasticity lower than that of the tungsten carbide one, a tin sheet

between the load applicator and crown as stress breaker, or a stainless steel loading

piston with its end machined to a curvature equivalent to 40-50 mm diameter to

reproduce clinically-relevant contact pressure. However, the increase in the loading

stresses of the tungsten ball would not have significantly affect the comparison between

the different test groups since all specimens were loaded using similar contact pressure.

Fracture load data reported in this study should be considered as relative, not absolute

values, and extrapolation of these in-vitro strength data to clinical performance must be

considered cautiously and within the limitations of the study.2

113

Conclusions

The veneering material, but not the core design, had a significant effect on fracture

load and mode of zirconia-based crowns. Lithium disilicate CAD-CAM-produced

veneers for zirconia copings seem to be a promising option for improving the in-vitro

performance of the crowns in comparison to feldspathic CAD-CAM-produced veneers.

114

Acknowledgements:


Toronto.

The work was supported by grants from the Faculty of Dentistry Research

Committee, University of Toronto. Materials and equipment donations were received

from Ivoclar-Vivadent, Patterson Dental Canada and Kuraray; and technical support

provided by Select Dental Lab (Woodbridge, ON, Canada) to whom the authors are

grateful.

115

(Table 5.1) Composition and some properties of materials used for core and veneering layer porcelain

(as reported by manufacturer)


Technique

CTE

(10-6 .K-1)

Flexure strength

(MPa)

Vita Mark II

(Lots: 17510 and

25520)

VitaZahnfabrik,

BadSackingen,

Germany

Feldspathic

porcelain CAD-CAM

9.4 0.1

(100 - 500 C) 154 15

IPS e.max CAD

(Lot: P14576)

Ivoclar Vivadent,

Schaan, Lichtenstein Lithium Disilicate CAD-CAM

10.5 0.4

(100 - 500 C) 360 60

IPS e.max ZirCad

(Lot: N77313)

Ivoclar Vivadent,


Yttrium-stabilized


10.75 0.25

(100 -400 C) 900 50

IPS e.max CAD

Crystall./Connect

(Lot: N65484 and

N69897)

Ivoclar Vivadent,


Fluorapatite glass-

ceramic

Manual

application

9.5 0.5

(100 -400 C) 160 620

Vita Glaze-LT

(Lot: 17810)

VitaZahnfabrik,

BadSackingen,

Germany

-

Manual

application - -

IPS e.max CAD

Crystall./glaze

Ivoclar Vivadent,


-

Manual

application

9.5 0.5

(100 -400 C) -

116

(Table 5.2) Firing programs for materials used in the study

Step

Beginning

temp.

(ºC)

Hold

time

(min.)

Temp. 1

increasing

rate

(ºC/min)

Temp.1

(ºC)

Hold

time 1

(min)

Temp. 2

increasing

rate

(ºC/min)

Temp. 2

(ºC)

Hold

time 2

(min)

Vacuum

On/Off

(ºC)

IPS e.max CAD-on

technique

Fusion/Crystallization

firing

403

6 (pre-

drying)

+ 2

30 820 2 30 840 7 550/840

IPS e.max CAD

Crystall./glaze 403 6 60 820 10 30 840 3 550/840

Vita Glaze-LT 500 4 80 780 1 80 - - -

117

(Table 5.3) Fracture load (N) and mode for all groups

Group

Fracture Load (N) Fracture Mode

N (%)

Mean (SD) Median

95 %

Confidence

Interval

Min - Max

Veneer

Only

Veneer

and Core

VMII non-anatomical cores 1796 (429) 1780 1489 - 2102 1202 - 2524 10 (100%) 0

VMII anatomical cores 2237 (323) 2278 2005 – 2468 1709 – 2748 10 (100%) 0

CAD non-anatomical cores 3427 (799) 3709 2856 - 3999 1734 - 4252 2 (20 %) 8 (80%)

CAD anatomical cores 3662 (1400) 3492 2660 - 4663 1097 – 5638 4 (40 %) 6 (60%)

118

(Table 5.4) Two-way ANOVA

(dependent variable: fracture load; independent variables: veneering technique and core design)

R Squared = 0.486 (Adjusted R Squared = 0.444)

Source Type III Sum

of Squares df Mean Square F Sig.

Corrected Model 24610699.28 3 8203566.427 11.367 < 0.001

Intercept 309213872.7 1 309213872.7 428.457 < 0.001

Veneering Technique 23363189.11 1 23363189.11 32.373 < 0.001

Core Design 1140728.156 1 1140728.156 1.581 0.217

Veneering Technique * Core

Design 106782.012 1 106782.012 0.148 0.703

Error 25980911.08 36 721691.975

Total 359805483.1 40


119

(Table 5.5) Fracture load according to fracture mode for EMAX crowns (N).

Core Design Non-Anatomical Anatomical

Fracture Mode Veneer Only Veneer and

Core Veneer Only

Veneer and

Core

1734 4252 4534 4451

3944 3754 2045 5328

3664 1097 5638

3417 3349 3433

2819 3194

3852 3550

4199

2639

Mean (SD) 2839 (1563) 3574 (590) 2756 (1502) 4266 (1039)

120

(MD: Mesio-distal and BL: bucco-lingual).

(Figure 5.1) Preparation dimensions of the master die (buccal view)

121

(Figure 5.2) Crown fabrication

a) Different core designs (mesial view).

b) Different core designs (buccal view).

c) Different core designs (occlusal view).

d) VMII veneer and respective coping after glazing.

e) EMAX veneer and respective coping before fusion and crystallization.

f) EMAX crown after glazing.

122

(Zr: zirconia, V: veneer and D: epoxy die)

(Figure 5.3) Fracture modes

a) All the VMII crowns showed fractures that involved the veneering layer only showing a cone crack

under the loading area in one of the VMII crowns.

b) 30% of the EMAX crowns showed fractures that involved the veneering layer only.

c) 70% of the EMAX crowns showed fractures that involved both the veneer and the core.

123

(Zr: zirconia and V: veneer)

(Figure 5.4) SEM picture for EMAX crown with non-anatomical core showing some flaws (arrow) between

the core material and the fusion ceramic

V

Zr

124

(Figure 5.5) SEM photomontage showing a fractured EMAX crown with anatomical core (fracture involved the veneering and core layers).

The fractographic examination revealed that the fracture started occlusally at the loading area with a cone crack and propagated through the veneering

layer. Crushing of the veneer ceramic can be observed at the loading area. (Zr: zirconia and V: veneer)

125

References




2004;48(2):viii, 513-viii, 30.

3. Anusavice KJ, Phillips RW. Phillips' science of dental materials. 1996;10th:66.


2008;24(3):299-307.



J Dent 2004;32(2):91-9.






2005;94(2):125-31.



2012;107(3):170-7.








2009;37(9):731-6.












126




Dent Mater 2009;25(7):857-62.



Prosthodont 2007;20(4):383-8.




2006;96(4):237-44.



Mater 2005;21(10):984-91.




22. Marchack BW, Futatsuki Y, Marchack CB, White SN. Customization of milled

zirconia copings for all-ceramic crowns: a clinical report. J Prosthet Dent

2008;99(3):169-73.



Prosthodont 2008;21(3):237-40.

24. Beuer F, Schweiger J, Eichberger M, Kappert HF, Gernet W, Edelhoff D. High-

strength CAD/CAM-fabricated veneering material sintered to zirconia copings--a

new fabrication mode for all-ceramic restorations. Dent Mater 2009;25(1):121-8.

25. Drummond JL, Eliades G. Ceramic behavior under different environmental and

loading conditions. Dental materials in vivo : aging and related phenomena:

Chicago ; London : Quintessence Pub. Co.; 2003. p. 35-45.

26. Rosenstiel SF, Land MF, Fujimoto J. Contemporary fixed prosthodontics. St.

Louis: Mosby; 2001. p. 643-72.

27. Evans AG. Slow crack growth in brittle materials under dynamic loading

conditions. International Journal of Fracture 1974;10(2):251-59.

28. Evans AG, Johnson H. Fracture stress and its dependence on slow crack growth.

Journal of Materials Science 1975;10(2):214-22.

29. Wiederho SM. Moisture assisted crack growth in ceramics. International Journal

of Fracture Mechanics 1968;4(2):171-77.

30. Choi JE, Waddell JN, Torr B, Swain MV. Pressed ceramics onto zirconia. Part 1:

Comparison of crystalline phases present, adhesion to a zirconia system and

flexural strength. Dent Mater 2011;27(12):1204-12.

127

31. Stawarczyk B, Ozcan M, Roos M, Trottmann A, Sailer I, Hammerle CH. Load-

bearing capacity and failure types of anterior zirconia crowns veneered with

overpressing and layering techniques. Dent Mater 2011;27(10):1045-53.

32. Lee JJ, Wang Y, Lloyd IK, Lawn BR. Joining veneers to ceramic cores and

dentition with adhesive interlayers. J Dent Res 2007;86(8):745-8.

33. Saied MA, Lloyd IK, Haller WK, Lawn BR. Joining dental ceramic layers with

glass. Dent Mater 2011;27(10):1011-6.

34. Guess PC, Bonfante EA, Silva NR, Coelho PG, Thompson VP. Effect of core

design and veneering technique on damage and reliability of Y-TZP-supported

crowns. Dent Mater 2013;29(3):307-16.

35. Kokubo Y, Tsumita M, Kano T, Fukushima S. The influence of zirconia coping

designs on the fracture load of all-ceramic molar crowns. Dent Mater J

2011;30(3):281-5.

36. Larsson C, El Madhoun S, Wennerberg A, Vult von Steyern P. Fracture strength

of yttria-stabilized tetragonal zirconia polycrystals crowns with different design:

an in vitro study. Clin Oral Implants Res 2012;23(7):820-6.



Dent 2009;37(12):978-83.



and veneer ceramics. J.Oral Rehabil. 2004;31(7):682-88.

39. Molin MK, Karlsson SL. Five-year clinical prospective evaluation of zirconia-

based Denzir 3-unit FPDs. Int J Prosthodont 2008;21(3):223-7.

40. Tinschert J, Schulze KA, Natt G, Latzke P, Heussen N, Spiekermann H. Clinical

behavior of zirconia-based fixed partial dentures made of DC-Zirkon: 3-year

results. Int J Prosthodont 2008;21(3):217-22.





43. Pallis K, Griggs JA, Woody RD, Guillen GE, Miller AW. Fracture resistance of

three all-ceramic restorative systems for posterior applications. J Prosthet Dent

2004;91(6):561-9.

44. Choi JE, Waddell JN, Swain MV. Pressed ceramics onto zirconia. Part 2:

indentation fracture and influence of cooling rate on residual stresses. Dent Mater

2011;27(11):1111-8.

45. Zahran M, El-Mowafy O, Tam L, Watson PA, Finer Y. Fracture Strength and

Fatigue Resistance of All-Ceramic Molar Crowns Manufactured with CAD/CAM

Technology. J Prosthodont 2008.

128




prosthodontic implications. Int J Prosthodont. 1995;8(2):105-16.

48. Alhasanyah A, Vaidyanathan TK, Flinton RJ. Effect of Core Thickness

Differences on Post-Fatigue Indentation Fracture Resistance of Veneered

Zirconia Crowns. J Prosthodont 2013.

49. Bhowmick S, Melendez-Martinez JJ, Zhang Y, Lawn BR. Design maps for

failure of all-ceramic layer structures in concentrated cyclic loading. Acta Mater

2007;55(7):2479-88.

50. Mainjot AK, Schajer GS, Vanheusden AJ, Sadoun MJ. Influence of cooling rate

on residual stress profile in veneering ceramic: measurement by hole-drilling.

Dent Mater 2011;27(9):906-14.

51. Hermann I, Bhowmick S, Lawn BR. Role of core support material in veneer

failure of brittle layer structures. J Biomed Mater Res B Appl Biomater

2007;82(1):115-21.





129


130

Characterization of the interface between zirconia and veneering porcelain

Authors:




Saudi Arabia


Professor and Head, Department of Restorative Dentistry, Faculty of Dentistry, University of Toronto,

Toronto, ON, Canada



Toronto, ON, Canada



Professor, Department of Restorative Dentistry, Faculty of Dentistry, University of Toronto, Toronto,

ON, Canada


Associate Professor and Head, Department of Biomaterials Science, Schulich School of Medicine &

Dentistry, The University of Western Ontario, London, ON, Canada






Phone: (416) 979-4934 X 4572

Fax: (416) 979-4936


This work was previously presented in part at the 39th General Session & Exhibition of the AADR,

Washington, DC, USA in 2010

131

Abstract

Objective: To evaluate the interface between the zirconium oxide core and veneering

porcelain using Time of Flight Secondary Ion Mass Spectrometry (ToF-SIMS) and X-ray

Photoelectron Spectroscopy (XPS). Methods: Four zirconium oxide slices (IPS e.max

ZirCad) were sintered to make rectangular specimens (1mm in thickness). A liner

material (IPS e.max Zirliner) was applied to two of the zirconia squares followed by

pressed (IPS e.max ZirPress) or manually applied (IPS e.max Ceram, Ivoclar Vivadent)

veneering porcelain. The other two squares were veneered as above but without the liner

material. The specimens were cross-sectioned using diamond discs on Leica TXP,

targeted sectioning and polishing unit (Leica Microsystems Inc). Before the surface

analysis, the samples were pre-cleaned with a 1 keV Ar+ ion beam. The chemical

composition changes along the interfacial areas of all the specimens (N=4) were

examined using Time of Flight Secondary Ion Mass Spectrometry (ToF-SIMS)

complemented by X-ray Photoelectron Spectroscopy (XPS). Results: Oxygen was the

main element present on both the veneer and core area. Sodium, silicon and aluminum

were the dominant elements in the veneering porcelain, while zirconium was the

dominant element in the zirconia core. XPS analyses revealed total disappearance of the

aluminum and sodium at the core region. A significant decrease (but not total

disappearance) in the relative atomic percentage of zirconium in the veneering porcelain

region was detected compared to the core region. ToF-SIMS data revealed variations in

zirconium chemistry in the interfacial region. Conclusions: The result of this study is

indicative of chemical changes of the zirconium along the interface regardless of the

veneering porcelain application technique or the usage of the liner. Further investigations

are needed to confirm and explain these changes.

132

Introduction

Several short-term clinical studies reported a high incidence of veneering porcelain

chipping.1-7

This high chipping rate, in comparison to that of metal-ceramic restorations,

led manufacturers and clinicians to seek explanations. Several hypothetical explanations

have been reported, including: occlusal overloading, residual stresses as a result of the

mismatch in the coefficient of thermal expansion (CTE) between core and veneering

porcelain, 8, 9

residual “tempering” stresses due to the thermal gradient between the

veneering porcelain and the zirconia core,9 interfacial stresses as a result of crystal

transformation at the interface,10


inherent strength

deficiency of the veneering porcelain 11, 12


Several studies have focused on measuring the zirconia-veneering porcelain bond

strength and comparing it to that of the metal-veneering porcelain.14-19

Other studies

focused on the effect of different factors on the zirconia-veneer bonding strength,

including: core-veneer CTE mismatch, use of the liner,20-22

surface treatment of the

core,21-27

veneering porcelain cooling rate28, 29

and veneering porcelain application

method and composition. 18, 25, 30-34

Unfortunately, little attention was directed to the

mechanism of the bonding between zirconia core and the veneering porcelain. It is not

clear whether the bond between the two materials is mechanical or a combination of both

chemical and mechanical. A recent study reported that zirconia and alumina as core

materials demonstrated mechanical bonding with silicate porcelain veneers, whereas

glass infiltrated alumina and lithium disilicate glass ceramic have reactive chemical

bonding towards silicate veneers. 35

The chemical structure at the core-veneer interface

was analyzed using Energy Dispersive X-ray (EDX), which revealed that some of the

veneer elements diffused into the zirconium oxide layer to a depth of 8 -10 μm.20

The objective of this study was to evaluate the nature of the bond between the

zirconium oxide core and matching veneering porcelain using the Time of Flight

Secondary Ion Mass Spectrometry (ToF-SIMS) and X-ray Photoelectron Spectroscopy

(XPS).

133

Materials and Methods:

Specimen preparation

Four zirconia rectangular specimens (11.5 x 12.5 x 1 mm) were prepared. These were

sliced from IPS e.max ZirCAD CEREC blocks (Size B40) (Ivoclar Vivadent, Schaan,

Lichtenstein) with a slow-speed diamond saw (Isomet, Buehler, Lake Buff, IL, USA).

The thickness of the slices was increased by 25% before cutting to compensate for the

shrinkage of zirconia that takes place during sintering. All slices were sintered in a high-

temperature furnace (Sintramat, Ivoclar Vivadent) according to manufacturer’s

instructions at a temperature of 1530◦C. The duration of the program, including the

cooling phase, was approximately 8 hours.

After sintering, a layer of ceramic bond (IPS e.max Zirliner, Ivoclar Vivadent)

was applied to one of the surfaces of two of the zirconia specimens, while the other two

samples were left without liner application. The Zirliner powder was mixed with build-

up liquid (Ivoclar Vivadent) to a creamy consistency, then layered onto the core

specimens, vibrated to achieve an even, greenish color, and finally fired in a porcelain

furnace (Whip Mix Pro 100, Whip Mix Corp, Louisville, KY, USA) according to the

manufacturer’s instructions (Table 6.1).

The zirconia specimens were then veneered with one of two veneering porcelains,

manually-applied or pressed, to produce four samples as follows: zirconia/manually-

applied veneer with liner (sample 1), zirconia/ manually-applied veneer without liner

(sample 2), zirconia/pressed veneer with liner (sample 3) and zirconia/pressed veneer

without liner (sample 4). The powder of the manually applied veneering material (IPS

e.max Ceram, Ivoclar Vivadent) was mixed with the manufacturer’s recommended liquid.

The obtained slurry was vibrated and condensed into the mold and the excess liquid was

removed by means of absorbent paper. The mold was removed and the specimens were

fired in a porcelain furnace (Whip Mix Pro 100) according to the manufacturer’s

instructions (Table 6.1). For the pressed veneer specimens, a wax pattern was applied on

the zirconia discs. The wax patterns were sprued and invested using a phosphate-bonded

investment material (IPS PressVEST Speed, Ivoclar Vivadent). The ceramic ingot (IPS

e.max Zirpress, Ivoclar Vivadent) was pressed according to the manufacturer’s

134

instructions using a pressing furnace (EP600 Combi, Ivoclar Vivadent) (Table 6.1). The

samples were sectioned perpendicular to the interface using a diamond saw (Leica TXP,

Leica Microsystems GmbH, Vienna, Austria).

ToF-SIMS analysis

Each sample was analyzed using ToF-SIMS IV, (IONTOF GMBH., Munster,

Germany). The area chosen for analysis was initial 191 x 191 μm2

straddling the

interfacial area for sample 1 and 500 x 500 μm2

for the other samples. First, The

analysis area was cleaned using a high current (90 nA, 3 keV) Ar+ ion beam for several

minutes until the main surface signal stabilized, indicating that the surface

contaminants had been reduced to steady state levels. A Bi+ ions (1.0 pA, 25 keV)

beam was used to generate the spectra in both high spatial and high spectral resolution

modes.36

Charge neutralization was achieved by using low-energy electron flooding.

Both positive and negative spectra were acquired and analyzed using

IonSpec/IonImage software (IONTOF GMBH). The collected spectra were used to

produce different ToF-SIMS pictures for the distribution of different ions in the

analysis area.

Samples of the liner and manually-applied veneering porcelain powders were also

separately analyzed to get their baseline spectra. Baseline spectra for the pressed

veneering porcelain and zirconia core material were not collected due to technical

difficulties.

• XPS analysis

As for the ToF-SIMS, the analysis area was sputter-cleaned with a 3keV Ar+ ion

beam. All XPS measurements were obtained with a Thermo Scientific K-Alpha X-ray

Photoelectron Spectrometer (East Grinstead, UK), with monochromated Al K-alpha X-

rays. Data were acquired across the interface using 30 μm spot size. Charge

compensation was performed using a low-energy flood. Vacuum pressure was 10-7

mbar,

with the vast majority of residual pressure from Argon associated with operation of the

135

charge compensation source. The spectra that were used for determining elemental

compositions were acquired in high pass energy snapshot mode. XPS data was analyzed

with Avantage software (East Grinstead, UK). Surface elemental compositions were

calculated from background-subtracted peak areas derived from transmission function

corrected regional spectra. Sensitivity factors (Scofield), which were provided by the

manufacturer, were used to calculate the relative atomic percentages.

136

Results and Discussion:

The chemical compositions of the different materials used in this study as provided

by the manufacturer are summarized in Table 6.2.

ToF-SIMS images for the 4 samples are presented in Figures 6.1 to 6.4 (brighter

areas indicate greater signals than the darker ones). The layers of the sample (core,

veneer and liner (if applied) can be clearly differentiated in these images. In agreement

with the chemical composition provided by the manufacturer, sodium, silicon and

aluminum were the dominant elements appearing in the liner and veneer layers, while

zirconium was the dominant element in the zirconia core. An interesting finding is the

appearance of zirconium across the interface although that the ToF-SIMS baseline

spectra of the liner and veneering porcelain did not show any presence of zirconium

(Figure 6.5). It is difficult to make a conclusion about the source of the zirconium across

the interface and in the veneering layer. Plausible explanations are the diffusion of the

zirconium from the core or a contamination from the diamond saw during sectioning.

The presence of the zirconium along the interface was confirmed and quantified using

the XPS technique. XPS data showed a significant decrease (but not total disappearance)

in the relative atomic percentage of zirconium at the interface region. In contrast,

elements like silicon, which is one of the main constituents of the veneering material,

disappeared completely at the interface region. XPS data for the changes in the relative

atomic percentage across the interface for all the 4 samples can be seen in Figures 6.6 to

6.9.

Furthermore, ToF SIMS data revealed a change in the zirconium chemistry away

from and near to the interface (Figure 6.10). The fragmentation patterns of the Zr (around

90 mass/units) and ZrxOy (around 105 mass/units) peaks are inherently related to the

chemistry of the system. Taking spectra from pixels in the region near the interfacial

boundary and comparing them to spectra taken away from this boundary reflected

differences in the chemistry of zirconium. This difference in the relative intensities of the

Zr/ZrH and ZrO/ZrOH signals is reflective of a real change in the zirconia chemistry.

As stated in the introduction, the chemical changes at the core-veneer interface was

analyzed in a previous study using Energy Dispersive X-ray (EDX), which revealed that

137

some of the veneer elements diffused into the zirconium oxide layer to a depth of 8 -10

μm.20

Although It is very difficult to confirm this observation in the current study due to

the difficulty of locating an exact point representing the interface, it has to be emphasized

that elements in the veneering and liner layers disappeared at different distances across

the interfacial region. XPS data showed that element like the sodium and aluminum

disappeared first followed by silicon.

While looking at the interface in cross-section is one of the approaches to study the

zirconia-veneer interface, this approach is limited depth-wise by the spatial resolution of

the ToF-SIMS (about 70 nm) and the XPS (about 10 μm). An alternative approach will

be to prepare samples with very thin veneering porcelain (about 10 nm) and look at the

interface through the veneering layer. Unfortunately, preparing such a sample is not

practical. Another approach would be to look at both sides of the fractured samples that

were fractured adhesively through the interface. This approach may eliminate the

possibility of contaminating the surfaces during sectioning and provide some insights

about the interface.

138

Conclusions:

ToF-SIMS and XPS examinations showed the qualitative and quantitative chemical

changes at the interfacial region. An important finding is the presence of zirconium atom

across the interfacial region, although that the chemical composition of both the

veneering and liner powders did not reveal any percentage of the zirconium atoms in

their composition. These changes might be indicative of chemical bonding between the

core and the veneer.

139

Acknowledgements:


Toronto.

The work was supported by grants from the Faculty of Dentistry Research Committee,

University of Toronto. Special thanks go to Mr. Peter Brodersen from Surface Interface Ontario

for his effort in performing the interfacial analysis.

140

(Table 6.1) Firing program for materials used in the study.

Material

Pre-drying

temperature

(ºC)

Holding time

(min.)

Heating rate

(ºC/min)

Final firing

temperature

(ºC)

Holding time

(min)

Vacuum

1

(ºC)

Vacuum

2

(ºC)


IPS e.max Ceram 403 4 40 750 1 450 749

IPS e.max ZirPress (100g

ring) 700 - 60 910 15 500 910

141

(Table 6.2) The chemical compositions of the different materials used in this study as provided by the

manufacturer.

Material Composition (% wt)

IPS e.max Ceram

Silicon dioxide (SiO2) (60-65%)

Aluminium oxide (Al2O3) (8-12%)

Sodium oxide (Na2O) (6-9%)

Potassium oxide (K2O) (6-8%)

Zinc oxide (ZnO) (2-3%)

CaO + P2O5 +F (2-6%)

Other oxides (2-8.5%)

Pigments (0.1 -1.5%)

IPS e.max Zirliner





CaO + P2O5 +F (2.5-7.5%)

Other oxides (1.5-8 %)

Pigments (0.1 -3%)

IPS e.max ZirPress





Calcium oxide (CaO) (2-4%)

Zirconium oxide (Zr O2) (1.5-2.5%)

P2O5 (1-2%)

F (0.5-1%)

Other oxides (0-6%)

Pigments (0.2-0.9%)

IPS e.max ZirCAD

Zirconium dioxide (ZrO2) (87-95%)

Yttrium oxide (Y2O3) (4-6%)

Hafnium oxide (HfO2) (1-5%)


142

(Figure 6.1) ToF-SIMS images for sample 1 (manually-applied veneering porcelain with liner).

Image size (191 x 191 μm2)

(A) ZrO2- (red) signal overlaid with AlO-

(green).

(B) ZrO2- (purple) signal overlaid with AlO-

(green) and Cl- (blue).

(C) ZrO+ (red) signal overlaid with Ca+ (green). (D) ZrO+ (yellow) signal

(E) Zr+ (yellow) signal

Liner

Core Veneer

Liner

Core Veneer

Core Veneer

Core Veneer

Core Veneer

143

(Figure 6.2) Tof SIMS image for sample 2 (manually-applied veneering porcelain without liner)


Zr- (green) signal overlaid with Li- (red).

Core Veneer

144

(Figure 6.3) ToF-SIMS images for sample 3 (pressed veneering porcelain with liner).


(A) Zr- (green) signal overlaid with Li-

(brown).

(B) Na+ (yellow) signal

Core Veneer

Liner Liner

Core Veneer

145

(Figure 6.4) ToF-SIMS image for sample 4 (pressed veneering porcelain without liner)


Zr- (green) signal overlaid with Li- (red).

Core Veneer

146

(Figure 6.5) Positive ToF-SIMS spectra for the IPS e.max Ceram powder (right) and Zirliner powder (left).

LiCH3

NaAl

Si

K Ca

64Zn

mass / u10 20 30 40 50 60 70 80 90

5x10

0.5

1.0

1.5

2.0

2.5

Inte

nsity

Li

CH3

NaAl

Si

K Ca

64Zn

mass / u10 20 30 40 50 60 70 80 90

5x10

0.2

0.4

0.6

0.8

1.0

1.2

1.4

1.6

Inte

nsity

147

Zirconia Veneer

(Figure 6.6) Changes in the relative atomic percentage as recorded by the XPS method across the

interfacial region for sample 1 (manually-applied veneering porcelain with liner).

0

10

20

30

40

50

60

70

0 200 400 600 800 1000 1200 1400

Rel

ativ

e A

tom

ic P

erce

nta

ge

Distance (micron)

C

O

Si

Zr

Na

Al

148

Zirconia Veneer


interfacial region for sample 2 (manually-applied veneering porcelain without liner).

0

10

20

30

40

50

60

70

0 200 400 600 800 1000 1200 1400

Rel

ativ

e at

om

ic P

erce

nta

ge

Distance (micron)

C

O

Si

Zr

Na

Al

149

Zirconia Veneer


interfacial region for sample 3 (pressed veneering porcelain with liner).

0

10

20

30

40

50

60

70

0 200 400 600 800 1000 1200 1400

Rel

ativ

e at

om

ic p

erce

nta

ge

Distance (micron)

C

O

Si

Zr

Na

Al

150

Zirconia Veneer


interfacial region for sample 4 (pressed veneering porcelain without liner).

0

10

20

30

40

50

60

70

0 200 400 600 800 1000 1200 1400

Rel

ativ

e at

om

ic p

erce

nta

ge

Distance (micron)

C

O

Si

Zr

Na

Al

151

(Figure 6.10) Tof-SIMS spectra showing a comparison of the Zr and ZrO regions from pixels close to and far away from the interface in the core area.

Zr/ZrH signal

ZrO/ZrOH signal

152

References

1. Raigrodski AJ, Hillstead MB, Meng GK, Chung KH. Survival and complications of zirconia-based fixed dental prostheses: a systematic review. J Prosthet Dent 2012;107(3):170-7.

2. Cehreli MC, Kokat AM, Akca K. CAD/CAM Zirconia vs. slip-cast glass-infiltrated Alumina/Zirconia all-ceramic crowns: 2-year results of a randomized controlled clinical trial. J Appl Oral Sci 2009;17(1):49-55.

3. Heintze SD, Rousson V. Survival of zirconia- and metal-supported fixed dental prostheses: a systematic review. Int J Prosthodont 2010;23(6):493-502.

4. Ortorp A, Kihl ML, Carlsson GE. A 3-year retrospective and clinical follow-up study of zirconia single crowns performed in a private practice. J Dent 2009;37(9):731-6.

5. Hosseini M, Worsaae N, Schiodt M, Gotfredsen K. A 1-year randomised controlled trial comparing zirconia versus metal-ceramic implant supported single-tooth restorations. Eur J Oral Implantol 2011;4(4):347-61.

6. Rinke S, Schafer S, Roediger M. Complication rate of molar crowns: a practice-based clinical evaluation. Int J Comput Dent 2011;14(3):203-18.

7. Sagirkaya E, Arikan S, Sadik B, Kara C, Karasoy D, Cehreli M. A randomized, prospective, open-ended clinical trial of zirconia fixed partial dentures on teeth and implants: interim results. Int J Prosthodont 2012;25(3):221-31.

8. de Kler M, de Jager N, Meegdes M, van der Zel JM. Influence of thermal expansion mismatch and fatigue loading on phase changes in porcelain veneered Y-TZP zirconia discs. J Oral Rehabil 2007;34(11):841-7.

9. Swain MV. Unstable cracking (chipping) of veneering porcelain on all-ceramic dental crowns and fixed partial dentures. Acta Biomater 2009;5(5):1668-77.

10. Tholey MJ, Swain MV, Thiel N. SEM observations of porcelain Y-TZP interface. Dent Mater 2009;25(7):857-62.

11. Sailer I, Feher A, Filser F, Gauckler LJ, Luthy H, Hammerle CH. Five-year clinical results of zirconia frameworks for posterior fixed partial dentures. Int J Prosthodont 2007;20(4):383-8.

12. Raigrodski AJ, Chiche GJ, Potiket N, Hochstedler JL, Mohamed SE, Billiot S, et al. The efficacy of posterior three-unit zirconium-oxide-based ceramic fixed partial dental prostheses: a prospective clinical pilot study. J Prosthet Dent 2006;96(4):237-44.

13. Aboushelib MN, de Jager N, Kleverlaan CJ, Feilzer AJ. Microtensile bond strength of different components of core veneered all-ceramic restorations. Dent Mater 2005;21(10):984-91.

153

14. Guess PC, Kulis A, Witkowski S, Wolkewitz M, Zhang Y, Strub JR. Shear bond strengths between different zirconia cores and veneering ceramics and their susceptibility to thermocycling. Dent Mater 2008;24(11):1556-67.

15. Al-Dohan HM, Yaman P, Dennison JB, Razzoog ME, Lang BR. Shear strength of core-veneer interface in bi-layered ceramics. J Prosthet Dent 2004;91(4):349-55.

16. Saito A, Komine F, Blatz MB, Matsumura H. A comparison of bond strength of layered veneering porcelains to zirconia and metal. J Prosthet Dent 2010;104(4):247-57.

17. Choi B-K, Han J-S, Yang J-H, Lee J-B, Kim S-H. Shear bond strength of veneering porcelain to zirconia and metal cores. The Journal of Advanced Prosthodontics 2009;1(3):129.

18. Lopez-Molla MV, Martinez-Gonzalez MA, Manes-Ferrer JF, Amigo-Borras V, Bouazza-Juanes K. Bond strength evaluation of the veneering-ceramics bonds. Medicina Oral Patología Oral y Cirugia Bucal 2010:e919-e23.

19. Blatz MB, Bergler M, Ozer F, Holst S, Phark JH, Chiche GJ. Bond strength of different veneering ceramics to zirconia and their susceptibility to thermocycling. Am J Dent 2010;23(4):213-6.

20. Aboushelib MN, Kleverlaan CJ, Feilzer AJ. Microtensile bond strength of different components of core veneered all-ceramic restorations. Part II: Zirconia veneering ceramics. Dent Mater 2006;22(9):857-63.

21. Aboushelib MN, Kleverlaan CJ, Feilzer AJ. Effect of zirconia type on its bond strength with different veneer ceramics. J Prosthodont 2008;17(5):401-8.

22. Mosharraf R, Rismanchian M, Savabi O, Ashtiani AH. Influence of surface modification techniques on shear bond strength between different zirconia cores and veneering ceramics. The Journal of Advanced Prosthodontics 2011;3(4):221.

23. Evil D, Ozden N, Celik E. Surface Roughness Effect on the Bonding of Y-Tzp Ceramics. J Dent Res 2008;87 (Spec Iss B): Abstract 2307.

24. Hill TJ, Tysowsky G. Effect of Sandblasting on Bond Strength of Three Zirconia Porcelains, . J Dent Res 2007;86 (Spec Iss A): Abstract 2448.

25. Kunii J, Kuriyama S, Hotta Y, Goldammer C, Tamaki Y, Fujishima A, et al. Effect of Porcelain Composition and Surface Treatments to Zirconia Bonding. J Dent Res 2009;88, (Spec Iss A): Abstract 1865.

26. Queiroz JR, Benetti P, Massi M, Junior LN, Della Bona A. Effect of multiple firing and silica deposition on the zirconia-porcelain interfacial bond strength. Dent Mater 2012;28(7):763-8.

27. Teng J, Wang H, Liao Y, Liang X. Evaluation of a conditioning method to improve core-veneer bond strength of zirconia restorations. J Prosthet Dent 2012;107(6):380-7.

154

28. Komine F, Saito A, Kobayashi K, Koizuka M, Koizumi H, Matsumura H. Effect of cooling rate on shear bond strength of veneering porcelain to a zirconia ceramic material. J Oral Sci 2010;52(4):647-52.

29. Göstemeyer G, Jendras M, Dittmer MP, Bach F-W, Stiesch M, Kohorst P. Influence of cooling rate on zirconia/veneer interfacial adhesion. Acta Biomaterialia 2010;6(12):4532-38.

30. Aboushelib MN, Kleverlaan CJ, Feilzer AJ. Microtensile bond strength of different components of core veneered all-ceramic restorations. Part 3: double veneer technique. J Prosthodont 2008;17(1):9-13.

31. Gelman A, Belgler M, Saleh N, Atlas AM, Blatz MB. Comparative Bond strength of a Pressed Veneering Ceramic to Zirconia. J Dent Res 2009;88 (Spec Iss A): Abstract 1869.

32. Ishibe M, Raigrodski AJ, Flinn BD, Chung KH, Spiekerman C, Winter RR. Shear bond strengths of pressed and layered veneering ceramics to high-noble alloy and zirconia cores. J Prosthet Dent 2011;106(1):29-37.

33. Aboushelib MN, de Kler M, van der Zel JM, Feilzer AJ. Effect of veneering method on the fracture and bond strength of bilayered zirconia restorations. Int J Prosthodont 2008;21(3):237-40.

34. Choi JE, Waddell JN, Torr B, Swain MV. Pressed ceramics onto zirconia. Part 1: Comparison of crystalline phases present, adhesion to a zirconia system and flexural strength. Dent Mater 2011;27(12):1204-12.

35. Y. Liu, Z. Shen, H. Feng. Chemical Mismatch of Bilayered All-ceramic Dental Composites. J Dent Res 2009;88, (Spec Iss A): Abstract 1862.

36. Sodhi RNS. Time-of-flight secondary ion mass spectrometry (TOF-SIMS): versatility in chemical and imaging surface analysis. Analyst 2004;129(6):483-7.

155

Chapter 7: Summary, Clinical Relevance and

Limitations

156

This thesis aimed to analyze the relatively high chipping rate of the veneering

porcelain of zirconia-based molar crowns; and to study some of the factors that may

affect it in-vitro. This type of crown failure may arise due to weakness of the veneering

porcelain or a weak bond between the core and veneer. The main focus of this thesis was

directed to evaluate some of the approaches that can improve the strength of the

veneering layer as follows:

1. Approaches addressing the inherent strength deficiency of the veneering

porcelain.

Two approaches aiming to improve the strength of the veneering porcelain

layer were evaluated. One involved the use of pressed vs. manually-applied

veneering porcelain which was mechanically-tested using simple-geometry

zirconia-porcelain specimens (Chapter 3). The results of this part favored the

pressing technique over the manual-application with a significant effect on the

fracture load, but not on the fracture mode.

In the other approach, CAD-CAM technology was used to produce veneering

porcelain with the aim to reduce/eliminate the flaws and porosities inherent to

the manually-applied and pressed veneering porcelains (Chapter 5). In addition,

this technique enabled the use of stronger porcelain (e.g. lithium disilicate) as a

veneering layer. The results of this part were promising with significantly higher

mean fracture strength for crowns veneered with CAD-CAM-fabricated lithium

disilicate veneers in comparison to ones veneered with traditional feldspathic

porcelain. While fracture modes of the lithium disilicate veneer group were

essentially catastrophic and involved both core and veneer, they occurred at

significantly higher fracture loads.

2. Approaches focusing on improving design of the crown.

In this project, the focus was directed to several design parameters: total

thickness, core thickness and core design. The effect of total thickness of the

zirconia-porcelain combination was assessed using flat bilayer specimens

(Chapter 3). Increasing the total thickness of the combination from 1.5 to 2 mm

resulted in a significant increase in the compressive fracture load of the bilayer

157

specimens. Clinically, a 2-mm thick crown would be ideally limited to heavily-

restored teeth or for dental implant applications. However, it should be avoided

when dealing with young patients with large pulp horns where a thinner crown

would be more appropriate.

The effect of the core thickness was evaluated using flat bilayer specimens

(Chapter 3) as well as zirconia-based anatomically-shaped crowns (Chapter 4).

Increasing the core thickness resulted in an increase of the fracture load of the

bilayer specimens as well as the molar crowns. Based on this finding, increasing

core thickness is recommended to increase overall strength of molar crowns.

However, long-term clinical trials are needed to substantiate this finding. In

addition, the effect of increasing the core thickness relative to the veneering

porcelain on the esthetics of the crowns should be considered.

The effect of the core design (anatomical vs. even thickness) on the

performance of the zirconia-based crowns was studied for both pressed and

CAD-CAM-produced veneering materials (Chapters 4 and 5, respectively). Both

studies revealed no significant effect of the core design on the fracture load.

However, the anatomical core design reduced the extent of veneering porcelain

fracture for the pressed type. This reduction in the size of the fracture can have a

significant clinical implication. Smaller fractures might translate into more

feasible chairside repairs as compared to the larger and more involving ones.

Clinical trials are needed to evaluate the effect of core design on clinical

performance of the zirconia-based crowns in comparison to the standard even

thickness design.

In addition to the studies above, one further study was directed to analyze the nature

of the zirconia-veneer bond (Chapter 6). Results of this study indicated that some

chemical bonding existed between zirconia and the veneering porcelain. Further

investigations are needed to verify such chemical bonding using other analytical

techniques of fractured specimens.

158

Finally, It has to be emphasized that the results of this project must interpreted with

caution as follows:

The results should be limited to the zirconia system used in this project.

Extrapolating the results to other systems might be difficult due to the possible

differences in the composition, microstructure, and mechanical and physical

properties of the components of different systems.

Although molar crowns were tested in this study, the results might be

generalized to the premolar area due to the similarity in the loading conditions.

Generalizing the results to the anterior region of the mouth might be challenging

due to the differences in the loading conditions.

Caution should be exerted in generalizing the results to implant-supported

crowns due to differences in the modulus of elasticity of the supporting

abutment material (titanium, precious metal alloys or zirconia).

159

Appendices

160

Appendix 1

161

Mechanical Cyclic Fatigue Studies

Author No. of samples Description Size Frequency No. of

Cycles

Load

Komine 125

8(cycled)

8 (not cycled)

For 1st maxillary

molar ceramic

crowns using

chewing simulator.

6 mm 1.3 Hz 1,200,000 0-49 N

Attia 61

8 (cycled)

8 (not cycled)

For 1st maxillary

premolar ceramic

crowns using

chewing simulator

At crosshead speed

(1mm/min).

4 mm ceramic ball 1.2 Hz 600,000 0-49 N

Azer 126

10 (static dry)

10 (static wet)

25 (cycled wet)

For posterior 3

cusps ceramic

crowns in wet

condition.

8 mm metallic

sphere

It was performed at the same loading rate as flexural testing (2.0

mm/min) with the use of a staircase approach. To simulate chewing

forces, this procedure involved initial loading of

the crown at 80 kg, stopping, lowering the load to 30 kg, stopping,

and returning to 80 kg until completion of 1000 cycles or fracture

(failure) of the crown. In each subgroup, if any crown survived 1000

cycles, the initial load for the next crown was raised by 10 kg. If the

crown fractured before the completion of 1000 cycles, the initial load

for the following crown was reduced by 10 kg. A minimum of 25

specimens was necessary to conduct the test.

Drummond 127

(Varies between 10-

15)

For ceramic

specimens

Crosshead speed

(2mm/min)

The cyclic fatigue testing in both air and water was done at the same loading rate as the

flexure testing using a staircase approach of either 1000 cycles or specimen fracture. This

procedure would initially load the ceramic bar at 4.0 kg, stop, lower the load to 1.0 kg, stop,

return to 4.0 kg, and then continue to cycle between 1.0 and 4.0 kg until 1000 cycles or

failure of the ceramic bar. If any ceramic bar survived the 1000 cycles, the initial load for

the next ceramic bar (a new specimen, not the one that did not fracture) was raised by 0.5

kg. If the ceramic bar did not survive, but fractured before completing 1000 cycles, the

initial load for the next ceramic bar was lowered by 0.5 kg. ( for ceramic specimens)

162


Cycles

Load

Chen29

20 (cycled)

20 (not cycled)

For 1st maxillary

molar ceramic

crowns using

artificial simulator

At crosshead speed

(0.5mm/min)

8 mm steel ball 1 Hz 50,000 200 N

Ohyama64

Using biaxial

flexure strength at

crosshead speed (1

mm/min)

1.5 mm spherical 20 Hz 100,000 4.9 – (60% of static fracture load ) N

Sobrinho66

16 (cycled)

10 (No cycled)

For premolar

crown shape at

crosshead speed of

(1mm/min)

4 mm stainless

steel ball

1 Hz 10,000 20-300 N

Sturb67

5 (cycled)

5 (not cycled)

For anterior crowns

at crosshead speed

(2mm/min)

Loading cylinder

wider than the

anterior teeth

separated from the

tooth with 0.8 mm

tinfoil

1.7 Hz 1,200,000 2-49 N

Kern63

6 (cycled)

14(not cycled)

For anterior bridges 4.76 mm stainless

steel ball on 3

unit-anterior

bridges

1,250,000 2-22 N

Kelly128

23 (Ambient)

17 (Wet)

For molar crowns 20 Hz 1,000,000 Ambient condition

(at 0,200,400,600 and 800 N)

Wet condition

(Staircase between 100 and 600N with

100 N step size)

163


Cycles

Load

Suputtamongkol129

For ceramic

specimens using

four-points bending

test in air and in

deionized water.

4.8 mm-diameter

steel ball

3 Hz 100,000 Maximum load of 200 N.

Gergawi 130

at least 6 specimens For crowns

cemented on epoxy

dies.

8mm sphere 1.6 Hz (baseline,

5000,

15000,

50000).

450 N

Vailati131

2 Hz (n=7)

or

10 Hz (n=18)

For ceramic

core:veneer disks

cemented to dentin

analog bases.

3 mm diameter

piston (dentin

analog material)

2 Hz or 10

Hz

500,000 A staircase sensitivity protocol was

followed, using a “localizing” step size

of 50 N and a “testing” step size of 25

N.

Tyan132

(Not cycled = 10)

(cycled =10)

For ceramic crowns

cemented to

standardized

acrylic dies

10,000 During cycling, the load on each

crown varied between the selected

maximum value (40% minimum

fracture load) and a minimum value of

about 1/3 of the maximum load.

Deesri133

(No cycled = 20)

(cycled =20)

For CEREC

conventional

designed molar

crown supported

with titanium

prefabricated post

and composite core

to the newly

designed CEREC

endocrown using a

fatigue simulator

MUC 2

50,000 200 N

Vult von et al. 81

(No fatigue =10)

(Mechanical

For crown shapes

cemented with Zn

phosphate on a

2.5 mm stainless

steel ball

1 Hz 10,000 30-300 N

164


Cycles

Load

fatiguing=10)

(Thermomechanical

fatiguing =10)

resinous dies.

165

Appendix II

166

Studies about the Clinical Performance of All-ceramic Crowns

Primary

Author Material

No of

restorati

ons

No of

patients

Survival

rate Fractures Duration

Posterior

Vs Anterior Description

Lehner 134

IPS Empress

(Ivoclar

Vivadent,

Schaan,

Liechtenstein)

78

(41

anterior

and 37

posterior)

34 95 % 4 fractures 19.7 +/- 8.5

months

posterior has

better survival Modified USPHS criteria

Gemalmaz135

IPS Empress

(Ivoclar)

37

(21

anterior

and

16

posterior)

20

The Kaplan-

Meier

survival rate

= 94.6%

2 crowns

fractured

24.56

months

California Dental Association's

quality assessment system

Sorensen 136

IPS Empress

(Ivoclar) 75 98.7 % 1 molar fractured 3 years

167

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Haselton 137

In-Ceram

Alumina

(VitaZahnfab

rik,

BadSackinge

n, Germany)

80

(67%

anterior

single

crowns,

26%

posterior

single

crowns,

6%

anterior

implant

crowns,

and 1%

posterior

implant

crowns)

41

1 molar fractured

1 MR of premolar

chipped.

1 occlusal wear

4 years Modified USPHS criteria

McLaren 138

In-Ceram

Alumina

(VitaZahnfab

rik)

223 96 %

Core fracture

0.6% per year

Porcelain fracture

0.3% per year

3 years

Anterior

crowns tended

to have a

slightly higher

3-year survival

rate (98%) than

premolars or

molars (94%).

168

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Segal 139

In-Ceram

Alumina

(VitaZahnfab

rik)

546

(32.4%

anterior [n

= 177] and

67.6%

posterior

[n = 369])

99.1%

2 core fractures

3 veneer fractures

6 years

The overall

success rate

was 99.1% (n =

541). The

success to

failure

ratio for

anterior was

98.9% to

1.1% ; and the

posterior crown

ratio was

99.2% to 0.8%

luted with glass ionomer resin

hybrid cement

Bindl 140

In-Ceram

Spinell core

crowns

(VitaZahnfab

rik)

19 (4

premolars

and 15

molars)

21

The Kaplan-

Meier

survival

rate=100%

No failure

39 +/- 11

months

Only posterior

crown were

included in this

study

Modified USPHS criteria

For crowns generated using

cerec 2 machine In-Ceram

Alumina core

crowns

(VitaZahnfab

rik)

24 (2

premolars

and 22

molars)

The Kaplan-

Meier

survival rate

=92%

2 molar fractured

Fradeani 141

In-Ceram

Spinell core

crowns

(VitaZahnfab

rik)

40 13

The Kaplan-

Meier

survival rate

= 97.5%

1 crown fractured

22 to 60

months

(mean 50

months)

Only anterior

crowns

California Dental Association's


169

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Bindl 142

In-Ceram

Spinell core

crowns

(VitaZahnfab

rik)

18

24

The Kaplan-

Meier

survival

rate=91.7%

1 crown fractured

44.7 +/-

10.3 months

Only anterior

crowns



cerec 2 machine

Vita Mark II

(VitaZahnfabr

ik)

18

The Kaplan-

Meier survival

rate =94.4%

1 crown fractured

Otto 143

Vita Mark II

(VitaZahnfab

rik)

20 20

All of the

crowns rated

clinically

acceptable

No fracture 1 year



CEREC 3 machine.

Fradeani 141

Procera

Allceram

(Nobel

Biocare,

Goteborg,

Sweden)

205

(50

anterior

and 155

posterior)

The Kaplan-

Meier

survival rate

= 96.7%

23.52

months

The Kaplan-

Meier survival

rate =100% for

the anterior

crowns and

95.15% for the

posterior

crowns

Odman144

Procera

Allceram

(Nobel

Biocare)

87 50

CSR of

97.7% and

93.5% (5-10

years

respectively)

Six crowns (7%) 5-10 years California Dental Association's


170

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Oden145

Procera

Allceram

(Nobel

Biocare)

100

crowns

58 93 %

3 (fracture of

veneer and core)

2 (fracture of

veneer)

5 years California Dental Association's


Cehreli115

Core: In-

Ceram

Zirconia.

(VitaZahnfa

brik)

Veneer: No

details

15

20

93 %

1 tooth fractured

and the crown

chipped (veneer

only)

2 years All teeth were

posterior

California Dental

Association's quality

assessment system

No details were given about

the zirconia crowns design.

Core:

Cercon

Zirconia.

(Degudent

GmbH,

Hanau-

Wolfgang,

Germany)

Veneer: No

details

15 93 %

1 crown fractured

(both core and

veneer)

171

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Encke124

Full-coverage

Gold 101

224

At 6 months

(100%)

At 1 year

(94.8%)

At 2 years

(92.7%)

None

2 years All teeth were

posterior

Monolithic

KaVo Everest

HPC (KaVo

Dental GmbH,

Biberach,

Germany)

123

At 6 months

(98.3%)

At 1 year

(95.9%)

At 2 years

(90.5%)

5 crowns partially

fractured

Beuer116

Core: IPS

e.max

ZirCad

(Ivoclar)

Veneer:

IPS e.max

Ceram

(Ivoclar)

50

38 patients

received 18

FDPs and

50 single

crowns

100% No fractures 35±14mont

hs

No details were given about the

zirconia crowns design

172

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Ortorp117

Core:

Procera

Zirconia

(Nobel

Biocare)

Veneer:

Vita Lumin

(VitaZahnfa

brik) or

NobelRondo

Zirconia

(Nobel

Biocare)

216 (25%

drop-out)

169 (19%

drop-out)

5-yaer CSR =

88.3%

2 posterior

crowns (veneer

chipping-

repairable)

4 crowns

(veneer chipping-

need replacement)

(1 anterior and 3

posteiors).

No core fracture

5 years

78% were

posterior

crowns

Zirconia core with anatomical

form (min. core

thickness=0.5mm and

veneering thickness between

1.0 and 2.0 mm)

173

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Hosseini 118

All-ceramic

Core:

KaVo

Zirconia

(KaVo

Dental

GmbH) or

Procera

Zirconia

(Nobel

Biocare)

Veneer:

HeraCeram

Zirkonia

(Heraeus

Kulzer

GmbH,

Hanau,

Germany)

and IPS

e.max

Ceram

(Ivoclar)

38

(implant-

supported)

36 100 % 1 year

All implants

replaced

mandibular

premolars



174

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Metal-

ceramic

Core:

HeraNordic

or ORION

WX .

Veneer:

HeraCeram

(Heraeus

Kulzer

GmbH) and

IPS d.SIGN

(Ivoclar

Vivadent)

37

(implant-

supported)

97.4 %

1 crown:

veneering

porcelain chipped

(Repairable)

Rinke119

All-ceramic

Core:

Cercon

Zirconia

(Degudent)

Veneer:

Cercon

Ceram-Kiss

(Degudent)

50 (4 drop-

out) 49

Crown

survival

rate=100%

Veneering

porcelain

success rate

= 97.9%)

1 crown:

veneering

porcelain chipped

(Repairable)

18.2 ± 4.6

months

All crowns

were on molars

Zirconia core with

anatomical form (min. core

thickness=0.4mm)



175

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Metal-

ceramic

Core:

Degunorm

(Degudent)

Veneer:

Duceragold

Kiss

(Degudent)

42 (2 drop-

out)

Crown

survival

rate=100 %

Veneering

porcelain

success rate

= 97.6%)

1 crown:

veneering

porcelain chipped

(Repairable)

Sagirkaya120

Core: Cercon

(Degudent),

ZirkonZahn

(Zirkonzahn

GmbH,

Bruneck,

Italy), Lava

(3M ESPE, St

Paul, MN,

USA) and

Katana

(Noritake)

Veneer: No

details

107 (74

tooth-

supported

and 33

implant-

supported)

42

At 1 year

(99.1%)

At 3 years

(97%)

At 4 years

(95.6%)

1 crown (veneer

chipping-need

replacement)

2 crowns (veneer

and core fracture)

4 years

(Mean

survival

time = 46.3

± 0.7

months)

California Dental


assessment system


the zirconia crowns design

176

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Tartaglia121

Core: Zirite

(Keramo,

Tavernerio,

Como,

Italia)

Veneer:

Feldspathic

porcelain

(CZR,

Noritake

Kizai Co.

Ldt,

Nagoya,

Japan)

248 single

crowns

(202 tooth-

supported,

36

implant-

supported

142

CSR was

given for the

crowns and

FDP

combined (5-

year CSR=

98.2%)

None of the single

crowns fractured. 3 years

190 posteriors

and 38

anteriors

Crown design: The core

generally supports a uniform

thickness of veneering

ceramic, and there should be

a maximum of 2 mm of

unsupported porcelain.

177

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Poggio122

Core:

Biotech

(Biotech srl, Nerviano, Italy), Diadem (Ivoclar

Vivadent),

IPS e.max

ZirCad

(Ivoclar

Vivadent),

Lava

Zirconia

(3M ESPE),

Procera

Zirconia

(Nobel

Biocare) or

Wieland

(Wieland

Dental

GmbH

Pforzheim,

Germany).

Veneer: No

details

102 31 Not reported

2 crowns had minor chipping of the veneering porcelain

20.9 ± 13.6

months

51 posteriors

and 51

anteriors

California Dental


assessment system

Crown design: Thinned zirconia margin, triangular in section, with height ranging from 0.5 mm in the vestibular areas up to 3.0

mm in the interproximal areas.

178

Primary

Author Material

No of

restorati

ons

No of

patients

Survival


Posterior


Groten123

Core:

Cercon

Zirconia.

(Degudent).

Veneer:

Cercon

Ceram Kiss.

(Degudent)

71 (17

drop-out)

32 (7 drop-

out)

2-year

Kaplan-Meier

survival

rate=98%

5 fractures (2

need to be

replaced)

Mean

observation

time = 21

months

California Dental


assessment system

CSR= Cumulative Survival Rate

179

Appendix III

180

Bond Strength of Zirconium Oxide to Veneering Porcelain

Author Core Veneering

porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Al-Dohan et al.91

Procera AllZirkon

(Nobel Biocare,

Goteborg,

Sweden)

Noritake CZR

(Noritake Kizai Co.

Ldt, Nagoya, Japan)

SBS test using

semicircular

knife edge at

crosshead speed

(0.5 mm/min)

28.03 (5.03) 41.6 % 57.4 %

DC-Zircon Vita D

(VitaZahnfabrik,

BadSackingen,

Germany)

27.9 (4.79) 58.6 % 40.4 %

Metal (Lodestar) Noritake (Noritake) 30.16 (5.89) 56 % 43 %

Giordano et al.146

LAVA core (3M

ESPE, St Paul,

MN, USA)

VM9

(Dentine wash)

(VitaZahnfabrik)

SBS using a

chisel like device

at crosshead

speed (0.5

mm/min)

38.45 (8.6)

Not reported

LAVA Ceram (3M

ESPE) 19. 2 (5.1)

Noritake (Noritake

Kizai Co. Ldt) 16.4 ± 9.2

In-Ceram YZ

(VitaZahnfabrik)

VM9

(Dentine wash) 34.2 (6.6)

181


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

(VitaZahnfabrik)

Blatz et al. 96

LAVA Core

(3M ESPE)

LavaCeram (3M

ESPE)

SBS at crosshead

speed (1

mm/min)

14.3 (0.6)

TC 14.8 (0.8)

Failure mode was cohesive within the

veneering porcelain for all the samples.

Cerabien ZR

(Cerabien ZR)

18.1(0.6)

TC 20.3 (0.7)

GC Initial

(GC America, Alsip,

IL, USA)

16.7 (0.7)

TC 17.5 (0.6)

Metal (Stability) VitaOmega

(VitaZahnfabrik)

8.0 (0.6)

TC 6.1 (0.5)

Aboushelib et al. 85

Cercon Base

(Degudent,GmbH,

Hanau-Wolfgang,

Germany)

Cercon Ceram S

(Degudent,GmbH,

Hanau-Wolfgang,

Germany)

MTBS at

crosshead speed

(1 mm/min) 29.1 (13.7) > 90 %

Aboushelib et al. 84

Cercon Base

(Degudent)

Cercon Ceram S

(Degudent)

MTBS at

crosshead speed

(1 mm/min)

No liner 17.2 (4.1)

w/ liner 26.3 (8.6)

60 %

60 %

Cercon Ceram Express

(Degudent)

No liner 38.6 (6.4)

w/ liner 29.7 (8.9)

100 %

95 %

182


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Noble Rondo Dentine

(Nobel Biocare,

Goteborg,

Sweden)

No liner 41.1 (11.1)

w/ liner 30.8 (14.5)

100 %

100 %

Noble Rondo Shoulder

(Nobel Biocare) No liner 39.3 (9.6) 70 %

LAVA Dentine (3M

ESPE)

No liner 30.9 (7.2)

w/ liner 34.3 (7.0)

70 %

55 %

Sakura Interaction

(Elephant Dental,

Hoorn, The

Netherlands)

No liner 19.9 (9.2)

w/ liner 23.8 (7.8)

90 %

95%

Experimental

pressable

porcelain

MTBS at

crosshead speed

(1 mm/min) No liner 25.2 (7.4) 95 %

Hill et al.102

e.max ZirCAD

(Ivoclar Vivadent)

e.max Ceram (Ivoclar

Vivadent)

SBS at crosshead

speed (1

mm/min)

Control=24.6±3.7

Sandblasted= 28.6±8.3

Not reported

183


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Lava Ceram

(3M/ESPE)

Control= 21.2±3.4

Sandblasted=17.2±6.

Vita VM9

(VitaZahnfabrik)

Control=18.0±5.9

Sandblasted=13.8±13.5

Evli et al.101

LAVA Core

(3M/ESPE)

Lava Ceram

(3M/ESPE)

SBS at crosshead

speed (0.5

mm/min)

Control = 279.77N

Sandblasted=265.00N

Plasma-

treated=201.77N

Wet grinding=272.30N

Not reported

Kunii et al. 97

Katana (Noritake)

CerabienZR

(Noritake)

3-Point bending

using chwickerth

crack initiation

test (Bond

strength was

calculated

mathematically)

Control= 28.6 ±2.3

Sandblasted= 26.9±1.7

Grinded= 27.6±1.5

Polished=26.3 ±2.1

Not reported IPS e.max Ceram

(Ivoclar vivadent) Control= 20.0±0.4

Zirox (Wieland) Control= 13.8±1.1

184


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Gelman et al.106

Katana (Noritake)

Nobel Rondo

(Nobel Biocare)

SBS at crosshead

speed (1

mm/min)

30.4±7.1

Not reported

VM9 (VitaZahnfabrik)

26.6±10.8

CZR (Noritake) 28.7±7.2

CZRpress

(Noritake)

29.5±4.2.

Guess et al. 94

Cercon Base

(Degudent)

Cercon CeramS

(Degudent)

SBS at crosshead

speed (5

mm/min) using

Schmitz-

Schulmeyer

test method)

Control=9.4±3.2

TC=9.6±4.2

Combined fracture modes: cohesive in the

veneer and adhesive at the core-veneer

interface.

Vita In-Ceram YZ

(Degudent)

VM9 (VitaZahnfabrik)

Control=12.5±3.2

TC=9.7±4.2

Predominant adhesive fractures at the core-

veneer interface.

DC-Zirkon

IPS e.max Ceram

(Ivoclar vivadent)

Control=11.5±3.4

TC=11.5±1.7

Predominant adhesive fractures at the core-

veneer interface.

185


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Degudent U94

(Degudent)

VM13 (Vita)

Control=27.6±12.1

TC= 26.4±13.4

Mainly cohesive fractures of the veneering

ceramic.

Aboushelib et

al.108

Not mentioned IPS e.max ZirPress

(Ivoclar vivadent)

MTBS

Crosshead speed

was not

mentioned.

26.6±1.6

Not reported

IPS e.max Ceram

(Ivoclar vivadent)

15.1±1.3

Aboushelib et al.99

Cercon Base

White

(Degudent)

Nobel Rondo

(Nobel Biocare)

MTBS

Crosshead speed

was not

mentioned.

As-milled=36.5±9.5

Abraded=42.4±11.5

Liner=28.5±15.3

65% Cohesive

100% Cohesive

80% Interfacial

Cercon Ceram

(Degudent)

As-milled=22.8±2

Abraded=37.9±5.1

Liner=31.6±7.7

60% Cohesive

100% Cohesive

65% Interfacial

186


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Cercon Base

Yellow

(Degudent)

Nobel Rondo

(Nobel Biocare)

As-milled=31.6±7.7

Abraded=24.3±8.7

Liner=29.3±7.3

60% Interfacial

85% Interfacial

90% Interfacial

Cercon Ceram

(Degudent)

As-milled=25.9±12.7

Abraded=17.2±5.3

Liner=37.1±12.2

55% Cohesive

70% Interfacial

70% Interfacial

LAVA Core

White

(3M/ESPE)

Nobel Rondo

(Nobel Biocare) As-milled=24.8±6.3

Abraded=29.7±7.3

Liner=23.4±11.4

90% Interfacial

80% Cohesive

70% Interfacial

Cercon Ceram

(Degudent)

As-milled=23±8.1

Abraded=36.1±8.1

Liner=41.8±9.4

90% Cohesive

70% Cohesive

70% Interfacial

LAVA Core

Yellow

(3M/ESPE)

Nobel Rondo

(Nobel Biocare)

As-milled=30.1±6.9

Abraded=20.8±10.3

Liner=29.4±7.4

80% Cohesive

100% Interfacial

70% Interfacial

187


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Cercon Ceram

(Degudent)


Abraded=16.8±5.2

Liner=39.3±7.7

70% Cohesive

80% Interfacial

75% Interfacial

Procera Zirconia

(Nobel Biocare)

Nobel Rondo

(Nobel Biocare)


Abraded=49.8±25.8

Liner=31.9±12.8

100% Interfacial

80% Cohesive

85% Interfacial

Cercon Ceram

(Degudent)

As-milled=33.9±5.6

Abraded=39.1±8.2

Liner=25.8±8.5

80% Cohesive

90% Cohesive

90% Interfacial

Aboushelib et

al.105

Cercon Base

(Degudent)


Posterior pressable

+

Cercon Ceram Kiss

(Degudent)

MTBS at

crosshead speed

(1 mm/min)

36.7 ± 5.1 72 %


Posterior pressable

(Degudent) +

Nobel Rondo Zirconia

(Nobel Biocare)

36.6 ± 5.1 77 %

188


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)


34.4 ± 2.9 100 %

Posterior pressable

(Degudent)

Aboushelib et

al.147

Cercon Base

(Degudent)

Nobel Rondo Press

Zirconia (Nobel

Biocare)

MTBS at

crosshead speed

(1 mm/min)

38.6 ± 6 Not reported

Nobel Rondo Zirconia

(Nobel Biocare) 39 ± 8

Choi et al. 93

Cercon Base

(Degudent) Cercon Ceram Kiss

(Degudent)

SBS at crosshead

speed (0.5

mm/min)

25.43 ± 3.12

Failure in zirconia group primarily occurred

near the interface with the residual veneering

porcelain remaining on the core (not

quantified)

Ni-Cr base metal

ceramic alloy

(Tillite, Talladium

Inc., LA,USA)

VM13

(Vita) 35.87 ± 4.23

High noble metal

ceramic alloy

(Degudent H,

Degudent)

VM13

(Vita) 38.00 ± 5.23

Fischer et al.111

In-ceram YZ

(Vita)

Allux (Wieland,

Pforzheim, Germany) SBS at crosshead

speed (1

mm/min)

27.7 ± 4.0 Failure mode was combination

(Interfacial/cohesive within the veneering

porcelain) for all the samples. Cerabien ZR 27.6 ± 3.9

189


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

(Noritake)

Creation ZI

(Metalordental,

Oensingen,

Switzerland)

24.4 ± 5.3

IPS e.max ceram

(Ivoclar)

23.5 ± 3.4

Initial ZR (GC, Tokyo,

Japan) 25.3 ± 4.8

Lava Ceram

(3M ESPE) 24.2 ± 5.6

Rondo Zirconia

(Nobel Biocare) 21.9 ± 6.2

Triceram

(Dentaurum,

Ispringen, Germany)

31.0 ± 7.1

Vintage ZR (Shofu,

Kyoto, Japan) 23.7 ± 5.3

VM9 (Vita) 30.8 ± 4.8

Zirox (Wieland) 22.6 ± 4.3

190


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Lopez-Molla et al. 95

CrNi alloy

(Ivoclar)

d.SIGN

(Ivoclar)

SBS at crosshead

speed (0.5

mm/min)

13.45

Not reported

IPS e.maxPress

(Ivoclar)

IPS e.maxCeram

(Ivoclar) 24.20

IPS e.maxZirCad

(Ivoclar)

IPS e.maxZirPress

(Ivoclar) 12.70

IPS e.maxZirCad

(Ivoclar)

IPS e.maxCeram

(Ivoclar) 7.86

Lava Frame

(3M ESPE)

LAVA Ceram

(3M ESPE) 10.20

Lava Frame

(3M ESPE)

IPS e.maxCeram

(Ivoclar) 4.62

Gostemeyer et al. 113

Lava Core

(3M ESPE)

LAVA

Ceram

(3M ESPE)

Rapid

cooling The bilayer

specimens were

notched; loaded

in a four-point

bending test and

8.2 ± 1.7 J/m2 Either the veneering ceramic was completely

delaminated with exposure of the zirconia

structure or a thin layer of the veneering

ceramic remained on the surface. Slow

cooling 7.5 ± 1.1 J/m2

191


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Triceram

(Dentauru

m)

Rapid

cooling

load-

displacement

curves were

recorded.

The critical load

to induce stable

crack extension

at the

core/veneer

interface was

determined, in

order to calculate

the strain energy

release rate.

13.3 ± 1.8 J/m2

Slow

cooling 9.8 ± 1.3 J/m2

VM9 (Vita)

Rapid

cooling 17.1 ± 2.8 J/m2

Slow

cooling 13.0 ± 1.9 J/m2

Zirox

(Wieland,

Pforzheim,

Germany)

Rapid

cooling 12.8 ± 2.0 J/m2

Slow

cooling 11.6 ± 1.5 J/m2

Mosharraf et al. 100

Cercon Base-

White (Degudent)

Cercon Ceram Kiss

(Degudent)

SBS at crosshead

speed

(1mm/min)

As milled=30.83 ± 5.79 100 %

Abraded=26.31 ± 2.53 90 % 10 %

Grinding=21.33 ± 4.22 100 %

Sandblasting +

Liner=27.39 ± 6.59 90 % 10 %

Cercon Base-

Colored

(Degudent)

As milled=25.73 ± 5.18 100 %

Abraded=28.51 ± 2.21 100 %

Grinding=25.39 ± 5.65 100 %

192


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Sandblasting +

Liner=29.50 ± 5.75 100 %

Komine et al.112

Katana (Noritake)

Super Porcelain AAA

(Noritake)

SBS at crosshead

speed

(0.5mm/min)

Rapid cooling=1.8 ±0.9

100 %

Slow cooling=

spontaneous debonding

CZR (Noritake)

Rapid cooling=27.5 ±

3.1 100 %

Slow cooling= 27 ± 4 100 %

IPS e.maxCeram

(Ivoclar)

Rapid cooling=22.1 ±

3.3 100 %

Slow cooling= 27.5 ±

2.7 100 %

Satio et al.92

Katana (Noritake)

CZR (Noritake)

SBS at crosshead

speed (0.5

mm/min)

27.0 ± 4.0

100 %

Cercon Ceram Kiss

(Degudent)

22.0 ± 1.3

100 %

IPS e.maxCeram

(Ivoclar)

22.1 ± 3.3

100 %

193


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Vintage ZR (Shofu) 24.8 ± 3.2

100 %

VM9 (Vita) 30.9 ± 3.8

100 %

Gold Alloy

(Degudent)

Super Porcelain AAA

(Noritake)

25.2 ± 5.7

12.5 % 87.5 %

Choi et al.98

In-Ceram YZ

(Vita)

CZR Press (Noritake)

The bilayer

specimens were

notched; loaded

in a four-point

bending test and

load-

displacement

curves were

recorded.

The critical load

to induce stable

crack extension

at the

core/veneer

interface was

determined, in

order to calculate

the strain energy

release rate.

26.7 ± 2.1 J/m2

Not reported

PM9 (Vita) 21.3 ± 2.07 J/m2

IPS e.max zirpress

(Ivoclar) 17.3 ± 3.2 J/m2

Press Xzr (Wieland) 17.2 ± 2.6 J/m2

194


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Ereifej et al. 148

IPS e.maxZirCad

(Ivoclar)

IPS e.maxCeram

(Ivoclar) SBS at crosshead

speed (0.2

mm/min)

28.8 ± 9.5

100 %

IPS e.max CAD

(Ivoclar)

IPS e.maxCeram

(Ivoclar) 29.1 ± 8.3 33 % 47 % 20 %

Ishibe et al.107

Olympia noble

Alloy (Jelenko, San Diego, CA,

USA)

IPS InLine POM (Ivoclar)

SBS at crosshead

speed (0.5

mm/min)

66.42 ± 20.60

Failure mode was cohesive within the

veneering porcelain for all the samples.

IPS InLine (Ivoclar) 63.69 ± 22.08

Pulse press-to-metal (Ceramay, Neu-Ulm,

Germany)

37.80 ± 20.57

Authentic Pulse (Ceramay)

43.74 ± 6.64

LAVA Core

(3M ESPE)

IPS e.maxZirPress

(Ivoclar)

40.41 ± 10.28

IPS e.maxCeram

(Ivoclar)

30.03 ± 9.49

PM9 (Vita) 21.34 ± 24.30

VM9 (Vita)

47.18 ± 12.99

195


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Teng et al.104

Cercon Base

(Degudent)

Cercon Ceram S

(Degudent)

SBS at crosshead

speed (0.5

mm/min)

Polished=36.66 ± 8.6

All specimens showed cohesive fractured in

the veneering porcelain +/- adhesive fracture

at the interface.

Abraded=39.14 ± 6.5

Modified with zirconia powder coating before sintering =47.02 ± 6.4

Nickel-chromium alloy (Argeloy N.P. (Be-Free); Argen Corp, San Diego, CA )

IPS Classic (Ivoclar) 46.12 ± 7.1

Queiroz et al.103

Lava core

(3M ESPE)

VM9 (Vita)

MTBS at

crosshead speed

(0.5 mm/min)

As milled (2 firing

cycles) = 15 ± 6.6

Failure initiated at the core–veneer interface

and propagated through the porcelain

As milled (3 firing

cycles) = 18.4 ± 6.1

Silica-coated (2 firing

cycles) = 13.8 ± 2.3

Silica-coated (3 firing

cycles) = 16.3 ± 2.9

196


porcelain Methods

Bonding Strength

Mean (SD) MPa

Type of Failure

Cohesive

(Veneer)

Interfacial

or combination

( Adhesive/

Cohesive)

Cohesive

(core)

Zeighami et al. 114

Cercon Base

(Degudent)

Cercon Ceram Kiss

(Degudent) (4 firing

cycles)

MTBS at

crosshead speed

(1 mm/min)

30.33 ± 2.13

95 % 5 %

Cercon Ceram Kiss


cycles)

27.43 ± 1.79 95 % 5 %

Cercon Ceram Kiss


cycles)

25.06 ± 1.76 90 % 5 % 5 %

SBS (Shear Bond Strength)

MTBS (Micro-Tensile Bond Strength)

TC (Thermocycled)

Abraded (Air-borne particle abrasion)

197

Reference

198

1. Rosenstiel SF, Land MF, Fujimoto J. Contemporary fixed prosthodontics. 3rd

ed.

St. Louis: Mosby; 2001. p. 262-71.

2. Shillingburg HT. Fundamentals of fixed prosthodontics. 3rd

ed. Chicago:

Quintessence Pub. Co; 1997. p. 433-54.

3. O'Brien WJ. Dental materials and their selection. 3rd

ed. Chicago: Quintessence

Pub. Co; 2002.


2004;48(2):viii, 513-viii, 30.


ed. St.

Louis, Saunders; 2003.

6. Sjogren G, Sletten G, Dahl JE. Cytotoxicity of dental alloys, metals, and ceramics

assessed by millipore filter, agar overlay, and MTT tests. J prosthet Dent

2000;84(2):229-36.

7. Kawai K, Urano M. Adherence of plaque components to different restorative

materials. Oper Dent 2001;26(4):396-400.

8. Yip KH, Smales RJ, Kaidonis JA. Differential wear of teeth and restorative

materials: clinical implications. Int J Prosthodont 2004;17(3):350-6.

9. Siervo S, Pampalone A, Siervo P, Siervo R. Where is the gap? Machinable

ceramic systems and conventional laboratory restorations at a glance.

Quintessence Int 1994;25(11):773-9.

10. Rosentritt M, Behr M, Lang R, Handel G. Influence of cement type on the

marginal adaptation of all-ceramic MOD inlays. Dent Mater 2004;20(5):463-69.

11. Kelly JR, Nishimura I, Campbell SD. Ceramics in dentistry: historical roots and

current perspectives. J prosthet Dent 1996;75(1):18-32.

12. Blatz MB. Long-term clinical success of all-ceramic posterior restorations.


13. Craig RG, Powers JM. Restorative dental materials. St. Louis: Mosby; 2002.

14. Drummond JL, Eliades G. Ceramic behavior under different environmental and

loading conditions. Dental materials in vivo : aging and related phenomena:

Chicago ; London : Quintessence Pub. Co.; 2003. p. 35-45.

15. Rosenstiel SF, Land MF, Fujimoto J. Contemporary fixed prosthodontics. 3rd

ed.

St. Louis: Mosby; 2001. p. 643-72.

16. Evans AG. Slow crack growth in brittle materials under dynamic loading

conditions. International Journal of Fracture 1974;10(2):251-9.

17. Evans AG, Johnson H. Fracture stress and its dependence on slow crack growth.

Journal of Materials Science 1975;10(2):214-22.

18. Bradt RC, Hasselman DPH, Lange FF. Fracture mechanics of ceramics: New

York ; London : Plenum Press; 1974. p. 613-46.

199

19. Wiederho SM. Moisture assisted crack growth in ceramics. International Journal

of Fracture Mechanics 1968;4(2):171-77.


ed.


21. Fatigue of Brittle Solids. In: Suresh S, editor. Fatigue of materials. Cambridge ;

New York: Cambridge University Press; 1991. p. xvii, 586 p.


restorations. J prosthet Dent 1999;81(6):652-61.

23. Anusavice KJ, Hojjatie B. Tensile stress in glass-ceramic crowns: effect of flaws

and cement voids. Int J Prosthodont 1992;5(4):351-8.

24. Morris DF, Kelly JR. Failure loads of bonded ceramics influenced by hydrostatic

compressive stresses. J Dent Res 1995;74:220.

25. Kon M, Kawano F, Asaoka K, Matsumoto N. Effect of leucite crystals on the

strength of glassy porcelain. Dent Mater J 1994;13(2):138-47.

26. Morena R, Lockwood PE, Fairhurst CW. Fracture toughness of commercial

dental porcelains. Dent Mater 1986;2(2):58-62.

27. Denry IL. Recent advances in ceramics for dentistry. Crit Rev Oral Biol Med.

1996;7(2):134-43.



J Dent 2004; 32(2):91-9.

29. Chen HY, Hickel R, Setcos JC, Kunzelmann KH. Effects of surface finish and

fatigue testing on the fracture strength of CAD-CAM and pressed-ceramic crowns.

J prosthet Dent 1999;82(4):468-75.

30. Giordano RA, Pelletier L, Campbell S, Pober R. Flexural strength of an infused

ceramic, glass ceramic, and feldspathic porcelain.J prosthet Dent 1995;73(5):411-

8.

31. Guazzato M, Albakry M, Quach L, Swain MV. Influence of grinding,

sandblasting, polishing and heat treatment on the flexural strength of a glass-

infiltrated alumina-reinforced dental ceramic. Biomaterials 2004;25(11):2153-60.

32. Kumin P, Luthy H, Mormann WH. The strength of ceramic and polymer after

CAD/CIM treatment and in a bond with dentin. Schweiz Monatsschr Zahnmed.

1993;103(10):1261-8.

33. Broderson SP. Complete-crown and partial-coverage tooth preparation designs

for bonded cast ceramic restorations. Quintessence Int 1994;25(8):535-9.

34. Hojjatie B, Anusavice KJ. Three-dimensional finite element analysis of glass-

ceramic dental crowns. J Biomech 1990;23(11):1157-66.




200


a layered zirconia and porcelain dental all-ceramic system.J prosthet Dent

2005;94(2):125-31.

37. Fleming GJ, El-Lakwah SF, Harris JJ, Marquis PM. The effect of core:dentin

thickness ratio on the bi-axial flexure strength and fracture mode and origin of

bilayered dental ceramic composites. Dent Mater 2005;21(2):164-71.


ed.


39. Isgro G, Pallav P, van der Zel JM, Feilzer AJ. The influence of the veneering

porcelain and different surface treatments on the biaxial flexural strength of a

heat-pressed ceramic.J prosthet Dent 2003;90(5):465-73.

40. Steiner PJ, Kelly JR, Giuseppetti AA. Compatibility of ceramic-ceramic systems

for fixed prosthodontics. Int J Prosthodont 1997;10(4):375-80.

41. Isgro G, Kleverlaan CJ, Wang H, Feilzer AJ. Thermal dimensional behavior of

dental ceramics. Biomaterials 2004;25(12):2447-53.

42. Isgro G, Kleverlaan CJ, Wang H, Feilzer AJ. The influence of multiple firing on

thermal contraction of ceramic materials used for the fabrication of layered all-

ceramic dental restorations. Dent Mater 2005;21(6):557-64.

43. Mackert JR, Evans AL. Quantitative X-Ray-Diffraction Determination of Leucite

Thermal-Instability in Dental Porcelain. Journal of the American Ceramic Society

1991;74(2):450-3.

44. Mackert JR. Effects of thermally induced changes on porcelain-metal

compatibility. Perspectives in dental ceramics: Quintessence Publishing, Chicago;

1985. p. 52-64.

45. Deng Y, Miranda P, Pajares A, Guiberteau F, Lawn BR. Fracture of

ceramic/ceramic/polymer trilayers for biomechanical applications. J Biomed

Mater Res A 2003;67(3):828-33.

46. Brukl CE, Philp GK. The fit of molded all-ceramic, twin foil, and conventional

ceramic crowns. J prosthet Dent 1987;58(4):408-13.

47. Tuntiprawon M, Wilson PR. The effect of cement thickness on the fracture

strength of all-ceramic crowns. Aust.Dent.J. 1995;40(1):17-21.

48. Scherrer SS, de Rijk WG. The fracture resistance of all-ceramic crowns on

supporting structures with different elastic moduli. Int J Prosthodont

1993;6(5):462-67.

49. Doyle MG, Goodacre CJ, Munoz CA, Andres CJ. The effect of tooth preparation

design on the breaking strength of Dicor crowns: 3. Int J Prosthodont

1990;3(4):327-40.

50. Burke FJ. Fracture resistance of teeth restored with dentin-bonded crowns: the

effect of increased tooth preparation. Quintessence Int 1996;27(2):115-21.

51. Julien KC, Buschang PH, Throckmorton GS, Dechow PC. Normal masticatory

performance in young adults and children. Arch Oral Biol 1996;41(1):69-75.

201


ed.



prosthodontic implications. Int J Prosthodont 1995;8(2):105-16.

54. McCormick JT, Rowland W, Shillingburg HT, Jr., Duncanson MG, Jr. Effect of

luting media on the compressive strengths of two types of all-ceramic crown.


55. Sobrinho LC, Cattell MJ, Knowles JC. Fracture strength of all-ceramic crowns. J

Mater Sci Mater Med 1998;9(10):555-9.

56. Rekow ED, Harsono M, Janal M, Thompson VP, Zhang G. Factorial analysis of

variables influencing stress in all-ceramic crowns. Dent Mater 2006;22(2):125-32.


ed.

Philadelphia: Saunders; 1996. p. 590-95.

58. Guazzato M, Albakry M, Ringer SP, Swain MV. Strength, fracture toughness and

microstructure of a selection of all-ceramic materials. Part II. Zirconia-based

dental ceramics. Dent Mater 2004;20(5):449-56.




60. Ellert DO, Kelly JR. All-Ceramic crown failure as a function of occlusal contact

location. J Dent Res 1997;76:392.

61. Attia A, Kern M. Influence of cyclic loading and luting agents on the fracture

load of two all-ceramic crown systems. J prosthet Dent 2004;92(6):551-6.

62. Kappert HF, Knode H, Schultheiss R. Strength of the In-Ceram system under

mechanical loads and thermocycling in artificial saliva. Dtsch Zahnarztl Z

1991;46(2):129-31.

63. Kern M, Douglas WH, Fechtig T, Strub JR, DeLong R. Fracture strength of all-

porcelain, resin-bonded bridges after testing in an artificial oral environment. J

Dent 1993;21(2):117-21.

64. Ohyama T, Yoshinari M, Oda Y. Effects of cyclic loading on the strength of all-

ceramic materials. Int J Prosthodont 1999;12(1):28-37.

65. Sobrinho LC, Cattell MJ, Glover RH, Knowles JC. Investigation of the dry and

wet fatigue properties of three all-ceramic crown systems. Int J Prosthodont

1998;11(3):255-62.

66. Sobrinho LC, Glover RH, Knowles JC, Cattell MJ. Comparison of the wet and

dry fatigue properties of all ceramic crowns. J.Mater.Sci.Mater.Med.

1998;9(9):517-21.

67. Strub JR, Beschnidt SM. Fracture strength of 5 different all-ceramic crown

systems. Int J Prosthodont 1998;11(6):602-09.

68. Draughn RA. Compressive fatigue limits of composite restorative materials. J

Dent Res 1979;58(3):1093-96.

202

69. Ritter J. Fracture Mechanics of Ceramics. NY: Plenum Publishing Co.; 1978.


fixed prosthodontics. J prosthet Dent 2003;90(1):31-41.

71. Piconi C, Maccauro G. Zirconia as a ceramic biomaterial. Biomaterials

1999;20(1):1-25.

72. Matsumoto RLK. Strength recovery in degraded yttria-doped tetragonal zirconia

polycrystals. J Am Ceram Soc 1985;86:C213.

73. Ishgi AA, Giordano R, Pober R. The effect of different surface treatments on

zirconia-ceramics' flexural strength. J Dent Res 2006; 85 (Spec Iss A): Abstract #

1870.

74. Sundh A, Molin M, Sjogren G. Fracture resistance of yttrium oxide partially-

stabilized zirconia all-ceramic bridges after veneering and mechanical fatigue

testing. Dent Mater 2005;21(5):476-82.

75. Papanagiotou HP, Morgano SM, Giordano RA, Pober R. In vitro evaluation of

low-temperature aging effects and finishing procedures on the flexural strength

and structural stability of Y-TZP dental ceramics. J Prosthet Dent

2006;96(3):154-64.

76. Raigrodski AJ. All-ceramic full-coverage restorations: concepts and guidelines

for material selection. Pract Proced Aesthet Dent 2005;17(4):249-56.

77. Sadan A, Blatz MB, Lang B. Clinical considerations for densely sintered alumina

and zirconia restorations: Part 1. Int J Periodontics Restorative Dent

2005;25(3):213-9.

78. Raigrodski AJ. Contemporary all-ceramic fixed partial dentures: a review. Dent

Clin North Am 2004;48(2):viii, 531-viii, 44.

79. Zahran M, El-Mowafy O, Tam L, Watson PA, Finer Y. Fracture Strength and

Fatigue Resistance of All-Ceramic Molar Crowns Manufactured with CAD/CAM

Technology. J Prosthodont 2008;17(5):370-7.



and veneer ceramics. J Oral Rehabil 2004;31(7):682-8.

81. Vult von SP. Fracture strength of two oxide ceramic crown materials after cyclic

preloading and thermocycling. Swed.Dent.J.Suppl 2005(173):1-21.

82. Wood KC, Berzins DW, Luo Q, Thompson GA, Toth JM, Nagy WW. Resistance

to fracture of two all-ceramic crown materials following endodontic access. J

Prosthet Dent 2006;95(1):33-41.

83. Thompson JY, Anusavice KJ, Naman A, Morris HF. Fracture surface

characterization of clinically failed all-ceramic crowns. J Dent Res

1994;73(12):1824-32.


different components of core veneered all-ceramic restorations. Part II: Zirconia

veneering ceramics. Dent Mater 2006;22(9):857-63.

203



Mater 2005;21(10):984-91.

86. Kelly JR, Tesk JA, Sorensen JA. Failure of all-ceramic fixed partial dentures in

vitro and in vivo: analysis and modeling. J Dent Res 1995;74(6):1253-8.

87. Ardlin BI. Transformation-toughened zirconia for dental inlays, crowns and

bridges: chemical stability and effect of low-temperature aging on flexural

strength and surface structure. Dent Mater 2002;18(8):590-5.

88. Chevalier J, Deville S, Munch E, Jullian R, Lair F. Critical effect of cubic phase

on aging in 3mol% yttria-stabilized zirconia ceramics for hip replacement

prosthesis. Biomaterials 2004;25(24):5539-45.

89. Guazzato M, Proos K, Quach L, Swain MV. Strength, reliability and mode of

fracture of bilayered porcelain/zirconia (Y-TZP) dental ceramics. Biomaterials

2004;25(20):5045-52.

90. Thompson GA. Influence of relative layer height and testing method on the

failure mode and origin in a bilayered dental ceramic composite. Dent Mater

2000;16(4):235-43.

91. Al-Dohan HM, Yaman P, Dennison JB, Razzoog ME, Lang BR. Shear strength

of core-veneer interface in bi-layered ceramics. J prosthet Dent 2004;91(4):349-

55.

92. Saito A, Komine F, Blatz MB, Matsumura H. A comparison of bond strength of

layered veneering porcelains to zirconia and metal. J prosthet Dent

2010;104(4):247-57.

93. Choi B-K, Han J-S, Yang J-H, Lee J-B, Kim S-H. Shear bond strength of

veneering porcelain to zirconia and metal cores. The Journal of Advanced

Prosthodontics 2009;1(3):129.




95. Lopez-Molla MV, Martinez-Gonzalez MA, Manes-Ferrer JF, Amigo-Borras V,

Bouazza-Juanes K. Bond strength evaluation of the veneering-ceramics bonds.

Medicina Oral Patología Oral y Cirugia Bucal 2010:e919-e23.

96. Blatz MB, Bergler M, Ozer F, Holst S, Phark JH, Chiche GJ. Bond strength of

different veneering ceramics to zirconia and their susceptibility to thermocycling.

Am J Dent 2010;23(4):213-6.

97. Kunii J, Kuriyama S, Hotta Y, Goldammer C, Tamaki Y, Fujishima A, et al.

Effect of Porcelain Composition and Surface Treatments to Zirconia Bonding. J

Dent Res 2009;88, (Spec Iss A): Abstract 1865.

98. Choi JE, Waddell JN, Torr B, Swain MV. Pressed ceramics onto zirconia. Part 1:

Comparison of crystalline phases present, adhesion to a zirconia system and

flexural strength. Dent Mater 2011;27(12):1204-12.

204

99. Aboushelib MN, Kleverlaan CJ, Feilzer AJ. Effect of zirconia type on its bond

strength with different veneer ceramics. J Prosthodont 2008;17(5):401-8.

100. Mosharraf R, Rismanchian M, Savabi O, Ashtiani AH. Influence of surface

modification techniques on shear bond strength between different zirconia cores

and veneering ceramics. The Journal of Advanced Prosthodontics 2011;3(4):221.

101. Evil D, Ozden N, Celik E. Surface Roughness Effect on the Bonding of Y-Tzp

Ceramics. J Dent Res 2008;87 (Spec Iss B): Abstract 2307.

102. Hill TJ, Tysowsky G. Effect of Sandblasting on Bond Strength of Three Zirconia

Porcelains, . J Dent Res 2007;86 (Spec Iss A): Abstract 2448.

103. Queiroz JR, Benetti P, Massi M, Junior LN, Della Bona A. Effect of multiple

firing and silica deposition on the zirconia-porcelain interfacial bond strength.

Dent Mater 2012;28(7):763-8.

104. Teng J, Wang H, Liao Y, Liang X. Evaluation of a conditioning method to

improve core-veneer bond strength of zirconia restorations. J Prosthet Dent

2012;107(6):380-7.


different components of core veneered all-ceramic restorations. Part 3: double

veneer technique. J Prosthodont 2008;17(1):9-13.

106. Gelman A, Belgler M, Saleh N, Atlas AM, Blatz MB. Comparative Bond

strength of a Pressed Veneering Ceramic to Zirconia. J Dent Res 2009;88 (Spec

Iss A): Abstract 1869.

107. Ishibe M, Raigrodski AJ, Flinn BD, Chung KH, Spiekerman C, Winter RR. Shear

bond strengths of pressed and layered veneering ceramics to high-noble alloy and

zirconia cores. J Prosthet Dent 2011;106(1):29-37.



Prosthodont 2008;21(3):237-40.

109. Y. Liu, Z. Shen, H. Feng. Chemical Mismatch of Bilayered All-ceramic Dental

Composites. J Dent Res 2009;88, (Spec Iss A): Abstract 1862.

110. M. Jakovac Akovac, J. Živiko-Babic, S. Musić, A. Carek. Contamination of

Zirconia Surface during Production in Dental Laboratory. J Dent Res 2009;88,

(Spec Iss A): Abstract 538.

111. Fischer J, Stawarzcyk B, Trottmann A, Hammerle CH. Impact of thermal misfit

on shear strength of veneering ceramic/zirconia composites. Dent Mater

2009;25(4):419-23.

112. Komine F, Saito A, Kobayashi K, Koizuka M, Koizumi H, Matsumura H. Effect

of cooling rate on shear bond strength of veneering porcelain to a zirconia

ceramic material. J Oral Sci 2010;52(4):647-52.

113. Göstemeyer G, Jendras M, Dittmer MP, Bach F-W, Stiesch M, Kohorst P.

Influence of cooling rate on zirconia/veneer interfacial adhesion. Acta

Biomaterialia 2010;6(12):4532-8.

205

114. Zeighami S, Mahgoli H, Farid F, Azari A. The Effect of Multiple Firings on

Microtensile Bond Strength of Core-Veneer Zirconia-Based All-Ceramic

Restorations. J Prosthodont 2013;22(1):49-53.




116. Beuer F, Edelhoff D, Gernet W, Sorensen JA. Three-year clinical prospective

evaluation of zirconia-based posterior fixed dental prostheses (FDPs). Clin Oral

Investig 2009;13(4):445-51.



2009;37(9):731-6.









121. Tartaglia GM, Sidoti E, Sforza C. A 3-year follow-up study of all-ceramic single

and multiple crowns performed in a private practice: a prospective case series.

Clinics (Sao Paulo) 2011;66(12):2063-70.

122. Poggio CE, Dosoli R, Ercoli C. A retrospective analysis of 102 zirconia single

crowns with knife-edge margins. J Prosthet Dent 2012;107(5):316-21.

123. Groten M, Huttig F. The performance of zirconium dioxide crowns: a clinical

follow-up. Int J Prosthodont 2010;23(5):429-31.

124. Encke BS, Heydecke G, Wolkewitz M, Strub JR. Results of a prospective

randomized controlled trial of posterior ZrSiO(4)-ceramic crowns. J Oral Rehabil

2009;36(3):226-35.

125. Komine F, Tomic M, Gerds T, Strub JR. Influence of different adhesive resin

cements on the fracture strength of aluminum oxide ceramic posterior crowns. J

Prosthet Dent 2004;92(4):359-64.

126. Azer SS, Drummond JL, Campbell SD, El Moneim ZA. Influence of core buildup

material on the fatigue strength of an all-ceramic crown. J Prosthet Dent

2001;86(6):624-31.

127. Drummond JL, King TJ, Bapna MS, Koperski RD. Mechanical property

evaluation of pressable restorative ceramics. Dent Mater 2000;16(3):226-33.



206

129. Suputtamongkol K, Anusavice KJ, Mecholsky JJ. Cyclic fatigue behavior of

baria-silicate glass-ceramics. J Dent Res 2002; 82 (Spec Iss A): Abstract #

1946(2005/04/17/).

130. Gergawi K, Giordano R, Pober R. Cyclic fatigue of CAD-CAM milled full

thickness and layered crowns. 2004; 83 (Spec Iss A): Abstract # 1367.

131. Vailati F, Kelly JR. Fatigue testing of cemented, bilayered, ceramic disks. J Dent

Res 2004; 83 (Spec Iss A): Abstract # 1366.

132. Tyan B, Pober R, Giordano R. Fatigue of CAD/CAM milled crowns. J Dent Res

2002; 82 (Spec Iss A): Abstract # 3963.

133. Deesri W, Kunzelmann KH, Ilie N, Hickel R. Fracture strength and weibull

evaluation of the Cerec endocrowns and post-and-core-supported conventional

Cerec crowns. J Dent Res 2003; 85 (Spec Iss B): Abstract # 0872.

134. Lehner C, Studer S, Brodbeck U, Scharer P. Short-term results of IPS-Empress

full-porcelain crowns. J Prosthodont 1997;6(1):20-30.

135. Gemalmaz D, Ergin S. Clinical evaluation of all-ceramic crowns. J Prosthet Dent

2002;87(2):189-96.

136. Sorensen JA, Choi C, Fanuscu MI, Mito WT. IPS Empress crown system: three-

year clinical trial results. J Calif Dent Assoc 1998;26(2):130-6.

137. Haselton DR, az-Arnold AM, Hillis SL. Clinical assessment of high-strength all-

ceramic crowns. J Prosthet Dent 2000;83(4):396-401.

138. McLaren EA, White SN. Survival of In-Ceram crowns in a private practice: a

prospective clinical trial. J prosthet Dent 2000;83(2):216-22.

139. Segal BS. Retrospective assessment of 546 all-ceramic anterior and posterior

crowns in a general practice. J Prosthet Dent 2001;85(6):544-50.

140. Bindl A, Mormann WH. An up to 5-year clinical evaluation of posterior in-ceram

CAD/CAM core crowns. Int J Prosthodont 2002;15(5):451-56.

141. Fradeani M, D'Amelio M, Redemagni M, Corrado M. Five-year follow-up with

Procera all-ceramic crowns. Quintessence Int 2005;36(2):105-13.

142. Bindl A, Mormann WH. Survival rate of mono-ceramic and ceramic-core

CAD/CAM-generated anterior crowns over 2-5 years. Eur J Oral Sci

2004;112(2):197-204.

143. Otto T. Computer-aided direct all-ceramic crowns: preliminary 1-year results of a

prospective clinical study. Int J Periodontics Restorative Dent 2004;24(5):446-55.

144. Odman P, Andersson B. Procera AllCeram crowns followed for 5 to 10.5 years: a

prospective clinical study. Int J Prosthodont 2001;14(6):504-9.

145. Oden A, Andersson M, Krystek-Ondracek I, Magnusson D. Five-year clinical

evaluation of Procera AllCeram crowns. J Prosthet Dent 1998;80(4):450-6.

146. Giordano R, Abrosa C, Ober R, Eucas R. Bond Strength of porcelain to Zirconia.

J Dent Res 2005; 84 (Spec Iss A): Abstract # 0544.

207

147. Aboushelib MN, de Kler M, van der Zel JM, Feilzer AJ. Microtensile Bond

Strength and Impact Energy of Fracture of CAD-Veneered Zirconia Restorations.

J Prosthodont 2009;18(3):211-6.

148. Ereifej N, Rodrigues FP, Silikas N, Watts DC. Experimental and FE shear-

bonding strength at core/veneer interfaces in bilayered ceramics. Dent Mater

2011;27(6):590-7.

factors affecting the fracture strength and fatigue ... · ii factors affecting the fracture...

Documents