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Aus der Universitätsklinik für Zahn-, Mund- und Kieferheilkunde der
Albert-Ludwigs-Universität Freiburg
Abteilung Poliklinik für Zahnärztliche Prothetik
(Ärztl. Direktor: Prof. Dr. J. R. Strub)
Fracture resistance of different Zirconia three-unit
posterior all-ceramic Fixed Partial Dentures
INAUGURAL-DISSERTATION zur Erlagung des
Zahnmedizinischen Doktorgrades
der Medizinischen Fakultät der Albert-Ludwigs-Universität Freiburg
Vorgelegt 2006
Von Kassiani Stamouli
Geboren in Aigio, Achaia, Griechenland
1
Dekan: Prof. Dr. Christoph Peters
1. Gutachter: Prof. Dr. J. R. Strub
2. Gutachter: Prof. Dr. J. Hausselt
Jahr der Promotion: 2006
2
Index
1 Introduction 6
2 Literature review 9
2.1 Ceramics 9 2.1.1 Historical perspectives of ceramics 9 2.1.2 Composition, properties and limitations of dental ceramics 10 2.1.3 Mechanisms of increasing the fracture resistance of ceramics: 11 2.1.4 Zirconia 14
2.1.4.1 Origin and applications 14 2.1.4.2 Properties 15 2.1.4.3 Zirconia products 16 2.1.4.4 Aging of zirconia ceramics 17
2.2 Classification of high-strength all-ceramic systems: 19 2.2.1 Glass-ceramics 19
2.2.1.1 Leucite Reinforced Glass-Ceramics 19 2.2.1.2 Lithium Disilicate Glass-Ceramics 20
2.2.2 Glass-infiltrated ceramics 20 2.2.2.1 In-Ceram Alumina® (Vita, D-Bad Säckingen) 20 2.2.2.2 In-Ceram Spinell® (Vita, D-Bad Säckingen) 21 2.2.2.3 In-Ceram Zirconia® (Vita, D-Bad Säckingen) 21
2.2.3 Polycrystalline ceramics 21 2.2.3.1 (CAD-)/CAM Systems: 21
2.2.3.1.1 Definition/Historical Background 21 2.2.3.1.2 CAD-CAM Components 22 2.2.3.1.3 Open/Closed (CAD-)/CAM systems 23 2.2.3.1.4 Materials 23 2.2.3.1.5 Yttrium Tetragonal Zirconia Polycrystals (Y-TZP) 24 2.2.3.1.6 The Cerec System 25 2.2.3.1.7 CAM Technologies 25 2.2.3.1.8 Marginal fit of CAD-CAM restorations 29
2.3 Clinical and technical aspects of all-ceramic FPDs 29 2.3.1 Clinical aspects 29
2.3.1.1 Preparation design 29 2.3.1.2 Translucency/Esthetics 29 2.3.1.3 Fracture resistance testing 30 2.3.1.4 Marginal Fit 32
3
2.3.1.5 Occlusal forces 32 2.3.1.6 Cementation 33
2.3.2 Technical aspects for all-ceramic FPDs 34 2.3.2.1 Connector dimensions 34 2.3.2.2 Thermal Expansion Coefficient (TEC) 35
2.4 Survival rates of all-ceramic FPDs 36
3 Aim of the study 39
4 Outline of the study (Fig. 4.1) 40
5 Materials and Methods 41
5.1 MATERIALS 41 5.1.1 Abutment Teeth 41 5.1.2 Materials used for the fabrication of the all-ceramic FPDs 41 5.1.3 Materials used for the cementation procedure 43 Impression and die materials 44 5.1.4 Additional Materials (Table 4): 45
5.2 METHODS 45 5.2.1 Representative model 45 5.2.2 Artificial periodontal membrane 46 5.2.3 Embedding models in the sample holders 46 5.2.4 Tooth preparation 46 5.2.5 Impression procedure 47 5.2.6 Fabrication of master models 47 5.2.7 Fabrication of all-ceramic FPDs 48
5.2.7.1 Manufacturing the framework 48 5.2.7.2 Veneering procedures: 50
5.2.8 Cementation of the FPDs 51 5.2.9 Dynamic loading of the test samples 52 5.2.10 Survival rate 55 5.2.11 Fracture resistance test 55 5.2.12 Statistics 55
6 Results 56
6.1 Survival rate of all-ceramic FPDs after aging 56
6.2 Fracture resistance tests 56
6.3 Fracture patterns 58
4
6.3.1 Procera Zirconia group 59 6.3.2 DCS group 60 6.3.3 Vita CerecInLab group 61
7 Discussion 62
7.1 Methods 62 7.1.1 The use of natural teeth as abutments 62 7.1.2 Artificial periodontal membrane 63 7.1.3 The antagonistic material 64 7.1.4 Preparation design and connector dimensions 64 7.1.5 Clinical relevance of fracture resistance tests 65 7.1.6 Clinical relevance of the artificial aging process 67
7.2 Results 68 7.2.1 Survival rate after the chewing simulation 68 7.2.2 Fracture resistance tests 68 7.2.3 Influence of the chewing simulation on the fracture resistance 69 7.2.4 Influence of the veneering process on the fracture resistance of zirconia-
based frameworks 70 7.2.5 Fracture patterns 71
8 Conclusions 73
9 Summary 74
10 Zusammenfassung 75
11 Appendix 76
11.1 Fracture resistance values of the Procera group 76
Without aging (Tab. 11.1) 76
11.2 Fracture resistance values of the DCS group 76
Without aging (Tab. 11.3) 76
11.3 Fracture resistance values of the Vita group 77
Without aging (Tab. 11.5) 77
12 References 78
13 Curriculum vitae 99
14 Acknowledgements 100
5
INTRODUCTION
1 Introduction
Most cultures throughout centuries have acknowledged teeth as an integral facial structure for
health, youth, beauty, and dignity. Unexpected loss of tooth structure and, particularly,
missing anterior teeth creates not only physical and functional problems, but often
psychological and social disturbances as well (Kelly et al, 1996). The treatment alternatives
for the replacement of a single missing tooth have expanded during recent times, so that the
choice of a proper treatment plan is no longer a simple decision.
A missing mandibular first molar is a relatively frequent dental problem. Treatment options to
replace a single missing molar include the Removable Partial Denture (RPD), the Fixed
Partial Denture (FPD), the Resin-Bonded Fixed Partial Denture (RBFPD) and the implant
supported crown (Priest, 1996a). In making the proper choice of the most appropriate
restoration type and material, one should consider both patient’s priorities and scientific
objectives (Priest, 1996a).
It is widely assumed that if posterior edentulous spaces are not treated, the adjacent teeth
ultimately will be lost. There is no sound scientific study, however, that describes the loss of
teeth associated with the inevitable disarrangement of one or both dental arches, triggered by
an unreplaced missing tooth (Shugars et al, 1998). According to the previous study, the vast
majority of untreated spaces did not result in loss of adjacent teeth within the study period.
Treatment with a removable partial denture did not increase the likelihood of adjacent tooth
survival, while treatment with a fixed partial denture resulted in an improved survival of
adjacent teeth.
In spite of their good survival rates of 86% after 5 years of function (Shugars et al, 1998),
84% and 59% after 5 and 10 years, respectively (Dietze, 2003), and low fabrication costs, the
indication of RPDs should be strictly limited because of high plaque accumulation, high risk
of caries and periodontitis progression and more frequent repair demand (Kerschbaum, 2004).
Therefore, replacing missing teeth by means of an RPD should only be applied when all other
treatment options would not be selected (Kerschbaum, 2004).
The RBFPD has gained in popularity over the years due to the rapid progress of adhesive
technology. The survival rates given in the literature are very divergent ranging from 65%
(maxillary region) and 40% (mandibular region) over 5 years (De Kanter et al, 1998) to 29%
over 6.2 years (Rijk et al, 1996) and 10% over 11 years of use (Priest, 1996b). The data
indicate that RBFPDs could be utilized as ideal interim restorations, offering a conservative,
quick and cost effective treatment option to the patient for a short or longer period of time.
6
INTRODUCTION
Different studies have reported survival rates ranging between 93-98% for implants and
between 83-98% for crowns with observation periods between 4 and 11 years, respectively in
partially edentulous cases (Scheller et al, 1998; Creugers et al, 2000; Naert et al, 2000;
Gibbard and Zarb, 2002; Mayer et al, 2002). The implant-supported single crown is
considered as a conservative (adjacent teeth remain intact), biocompatible and potentially
excellent esthetic treatment alternative. However, its application may be contraindicated due
to esthetic limitations, health considerations or negative patient compliance.
The traditional porcelain-fused–to-metal fixed partial denture (PFM) is the most popular
treatment option for the majority of dentists because of the familiar fabrication techniques, the
acceptable esthetic outcome and the high survival rates ranging between 74% and 85% after
15 years of service (Creugers et al, 1994; Scurria et al, 1998; Walton, 2002). The frequent
gingival discoloration around the metal margins of PFMs (Christensen, 1994), together with
some allergic reactions by metal alloys, are still a weak point of these restorations that
dissatisfies both patients and dentists (Shepard et al, 1983; Hansen and West, 1997). The
development of facial porcelain margins is one significant modification that enhances
esthetics by eliminating the display of metal and allowing a more natural transmission of light
(Hobo and Shillingburg, 1973; Shillingburg et al, 1973; Goodacre et al, 1977; Chiche et al,
1986). Using this technique, the framework is shortened by 1 to 3 mm in the shoulder area.
Other authors suggested that a minimum facial metal reduction of 2 mm is necessary in order
to obtain proper light transmission after cementation, which may compromise the fracture
strength (O'Boyle et al, 1997). It was further concluded that collarless metal-ceramic crowns
having up to 2 mm of unsupported porcelain could resist the same axial pressure as
restorations with complete metal strength, provided that a 90-degree shoulder tooth
preparation is used (Lehner et al, 1995).
Since the application of PFMs has proven to be successful over the years, these restorations
still remain the gold standard in terms of predictability. Despite this success, however, the
demands for more esthetic materials with biocompatible properties is increasing. The use of
metal in the oral cavity has come under dispute in recent years due to its eventual biological
incompatibility risks (Pfeiffer and Schwickerath, 1989; Reuling et al, 1990; Lucas and
Lemons, 1992; Rechmann, 1993). Therefore, all-ceramic restorations are considered as an
alternative of high importance and clinical value. After the introduction of feldspathic
porcelain reinforced with alumina (McLean and Hughes, 1965), researchers have been
developing new high-strength ceramic materials that can be used for the fabrication of FPDs
for use in both the anterior and posterior regions. In vitro and in vivo investigations of these
7
INTRODUCTION
newly-developed all-ceramic systems should be undertaken before introducing them into
routine clinical use.
The aim of the present study was to compare the fracture resistance and mode of failure of
different zirconia three-unit posterior all-ceramic fixed partial dentures and to evaluate the
effect of fatigue loading on the fracture resistance.
8
LITERATURE REVIEW
2 Literature review
2.1 Ceramics
2.1.1 Historical perspectives of ceramics
During the 18th century, the candidate materials for replacing teeth were human teeth, carved
animal teeth, ivory, and mineral or porcelain teeth. In 1723, Piere Fauchard was credited with
recognizing the potential of porcelain enamels and initiating research with porcelains to
imitate the color of teeth and gingival tissues (Jones, 1985).
In 1774, Alexis Duchateau and Nicholas Dubois de Chemant fabricated the first successful
porcelain dentures. Dubois de Chemant, who improved porcelain formulations continually
during his scientific career, was awarded both French and British patents.
In 1808, in Paris, Giuseppangelo Fonzi introduced individually-formed porcelain teeth that
contained embedded platinum pins. Their esthetic and mechanical versatility provided a major
advance in prosthetic dentistry.
As early as at the end of the 19th century, all-ceramic restorations, called jacket crowns, were
fabricated by firing a feldspathic ceramic material on a die prepared with platinum foil. Jacket
crowns were the only fixed esthetic restorations available at that time (Freese, 1959). Despite
their esthetic advantages, the restorations failed to gain widespread popularity because of their
high probability of fracture, low strength and poor marginal seal. This technique went out of
fashion once the metal-ceramic era began (Jones, 1985). A noteworthy development occurred
in the 1950s, with the addition of leucite to porcelain formulations that elevated the
coefficient of thermal expansion to allow their fusion to certain gold alloys to form complete
crowns and FPDs (Freese, 1959; Weinstein, 1962).
9
LITERATURE REVIEW
2.1.2 Composition, properties and limitations of dental ceramics
Dental ceramics consist of a compound of metals (aluminium, calcium, lithium, magnesium,
potassium, sodium, tin, titanium, and zirconium) and nonmetals (silicon, boron, fluorine, and
oxygen) that may be used as a single structural component, such as when used for a CAD-
CAM inlay, or as one of several layers used for the fabrication of a ceramic-based restoration.
Conventional dental porcelain is a vitreous ceramic based on a silica (SiO2) network and
potash feldspar (K2O·Al2O3·6SiO2), soda feldspar (Na2O·Al2O3·6SiO2) or both. Pigments,
opacifiers and glasses are added to control the fusion temperature, sintering temperature,
thermal contraction coefficient, and solubility. The feldspars used for dental porcelains are
relatively pure and colorless. Therefore, pigments must be added to produce the hues of
natural teeth (Anusavice, 2003). Most of the ceramics are characterized by their refractory
nature, hardness, and chemical inertness. A hardness of a ceramic similar to that of enamel is
desirable to minimize the wear of resulting ceramic restorations, and reduce the wear damage
that can be produced on enamel by the ceramic restoration. Chemical inertness ensures that
the surface of dental restorations does not release potentially harmful elements, and reduces
the risk for surface roughening and an increased susceptibility to bacterial adhesion to insure
excellent biocompatibility over time. Furthermore, ceramics demonstrate excellent insulating
properties, such as low thermal conductivity, low thermal diffusivity, and low electrical
conductivity. Their most attractive property is their potential for matching the appearance of
natural teeth, offering great esthetic results (Anusavice, 2003).
On the other hand, the susceptibility of ceramics to brittle fracture is a drawback, particularly
when flaws and tensile stresses coexist in the same region of the restoration. The flaw can be
a microcrack on the surface (e.g. created during occlusal adjustment with a diamond stone), or
it can be a subsurface porosity (e.g. from a processing error during the build-up and baking of
the porcelain) (Rosenblum and Schulman, 1997). When tension stress is applied, small flaws
tend to open up and propagate cracks (crack propagation theory) (O' Brien, 2002).
Irregularities in a bulk of the material, such as discontinuities and/or abrupt changes in shape
or thickness in the ceramic contour, act as stress raisers, making the restoration more prone to
failure. Stress around a stress raiser is higher than the average stress in the body of the
material. The amount of this increased stress depends on the shape of the stress raiser (e.g.�
stress at the tip of a sharp notch would be greater than that of a semicircular groove). Because
of the stress concentration at surface scratches and other defects (brittleness), ceramics tend to
fail at stress levels that are much lower than the theoretical strength to be tolerated. Compared
10
LITERATURE REVIEW
to metals, which can yield to high stress by deforming plastically, ceramics tend to have no
mechanism for yielding to stress without fracture (O' Brien, 2002). Therefore, cracks may
propagate through a ceramic material at low average stress levels. As a result, ceramics and
glasses have lower tensile strengths than compressive strengths (O' Brien, 2002).
2.1.3 Mechanisms of increasing the fracture resistance of ceramics:
1. Development of residual compressive stresses:
The thermal expansion coefficient (TEC) of the core ceramic is slightly greater than that of
the veneering ceramic. This mismatch allows the core material to contract slightly more upon
cooling from the firing temperature to room temperature, and leave the veneering ceramic in
residual compression while offering additional strength (Mackert, 1988).
2. Minimize the number of firing cycles:
Firing procedures sinter the particles densely together and produce a relatively smooth
surface. In addition, they increase the concentration of leucites in the porcelain, which in turn
leads to an increase of the TEC and a further mismatch between core/veneering porcelain.
This mismatch will cause immediate or delayed crack formation in the porcelain (Fairhurst et
al, 1980; Mackert, 1988; Mackert and Evans, 1991; Fairhurst et al, 1992).
3. Minimize tensile stress through optimal design of ceramic restorations
Dental restorations containing ceramics should be designed in a way to overcome their
weaknesses. The design should avoid exposure of the ceramic to high tensile stresses
(Anusavice, 2003). In the case of a crown, tensile stresses can be reduced by using strong core
materials with appropriate thickness, since these stresses are distributed on the inner surface
(core material is in tension) (Kelly et al, 1989; White et al, 1994; Zeng et al, 1996;
Wakabayashi and Anusavice, 2000; Lawn et al, 2001). In the case of a FPD, high tensile
stresses develop at the gingival surface of the connector� and a larger radius of curvature at
the gingival embrasure reduces the concentration of tensile stresses, thus affecting the fracture
resistance of the FPD (Oh et al, 2002; Oh and Anusavice, 2002). To promote achieving the
required connector dimensions without compromising the health of the supporting tissues, it
11
LITERATURE REVIEW
was suggested to fabricate the gingival and lingual aspects of the connectors exclusively out
of the framework material (McLaren, 1998).
4. Ion Exchange (or chemical tempering):
This process involves the exchange of larger potassium ions for the smaller sodium ions (a
common constituent of a variety of glasses)(Anusavice et al, 1992). If a sodium-containing
glass article is placed in a bath of molten potassium nitrate, potassium ions in the bath
exchange places with some of the sodium ions on the surface of the glass particles. The
potassium ion is about 35% larger than the sodium ion. Squeezing of the potassium ion into
the place formerly occupied by the sodium ion creates large residual compressive stresses in
the surfaces of the glasses subjected to this treatment. However, the depth of the compression
zone is less than 100 µm, so that this effect would be easily worn out after long–term
exposure to certain inorganic acids (Southan, 1970; Jones, 1983; Seghi et al, 1990; Anusavice
et al, 1992; Anusavice et al, 1994).
5. Thermal Tempering:
This is a process of creating residual surface compressive stresses by rapidly cooling the
surface of the object while it is hot and in the softened (molten) state. This rapid cooling
produces a skin of rigid glass surrounding a soft (molten) core. As the molten core solidifies�
it tends to shrink, but the outer skin remains rigid. The pull of the solidifying molten core, as
it shrinks, creates residual tensile stresses in the core and residual compressive stresses within
the outer surface, inhibiting the initiation and the growth of cracks (Anusavice et al, 1989;
Anusavice and Hojjatie, 1991; DeHoff and Anusavice, 1992)
6. Dispersion strengthening:
This involves the reinforcement of ceramics with a dispersed phase of a different material that
is capable of hindering a crack from propagating. Dental ceramics containing primarily a
glass phase can be strengthened by increasing the crystal content of leucite, lithium disilicate,
alumina, magnesia-alumina spinel, zirconia and other types of crystals (McLean and Hughes,
1965).
12
LITERATURE REVIEW
When a tough, crystalline material such as alumina (Al2O3) is added to a glass, the glass is
toughened and strengthened, because the crack cannot pass through the alumina particles as
easily as it can pass through the glass matrix (McLean and Hughes, 1965; Jones, 1983).
The amount of toughening depends on the crystal type, its size, its volume fraction, the
interparticle spacing, and its relative thermal expansion coefficient to the glass matrix. In
most instances, the use of a dispersed crystalline phase to disrupt crack propagation requires a
close match between the thermal contraction coefficients of the crystalline material and the
surrounding glass matrix (Jones, 1983).
7. Transformation toughening:
The dispersion strengthening process relies on the toughness of the particle to absorb energy
from the crack and deplete its driving force for propagation. The transformation toughening
process relies on a crystal structural change of a material under stress to absorb energy from
the crack (Morena, 1986). Zirconia (ZrO2) ceramic is a good example for this mechanism.
The material is polymorph occurring in three forms: monoclinic (M), tetragonal (T) and
cubic(C). Pure zirconia is monoclinic in room temperature. This phase is stable up to 1170°C.
Above this temperature it transforms into tetragonal and then into a cubic phase at 2370°C.
When ZrO2 is heated above 1170°C, the transformation from the monoclinic to the tetragonal
phase is associated with a 5% volume decrease. Reversely, during cooling, the transformation
from the tetragonal to the monoclinic phase is associated with a 3% volume expansion. These
phase transformations, however, induce stresses which result in crack formations. The
inhibition of these transformations can be achieved by adding stabilizing oxides (CaO, MgO,
Y2O3), which allow the existence of tetragonal-phase particles at room temperature. When
sufficient stress develops in the tetragonal structure and a crack in the area begins to
propagate, the tetragonal grains transform to monoclinic grains. The associated volume
expansion results in compressive stresses at the edge of the crack front and extra energy is
required for the crack to propagate further (Tateishi, 1987).
13
LITERATURE REVIEW
2.1.4 Zirconia
2.1.4.1 Origin and applications
Zirconia, the metal dioxide (ZrO2), was identified in 1789 by the German chemist Martin
Heinrich Klaproth in the reaction product obtained after heating some gems. It was used for a
long time, blended with rare earth oxides, as pigments for ceramics. The first biomedical
application of Zirconia, was carried out in 1969 by Helmer and Driskell (Helmer, 1969),
while the first use of zirconia in orthopedics was introduced by Christel (Christel, 1988) to
manufacture ball heads for total hip replacements. Its application over the years was further
expanded in dentistry; including the fabrication of brackets in orthodontics (Keith et al, 1994),
post and core systems (Edelhoff and Sorensen, 2002; Heydecke et al, 2002) and ceramic
implants/implant abutments offering improved esthetic alternatives (Glauser et al, 2004;
Kohal et al, 2004).
14
LITERATURE REVIEW
2.1.4.2 Properties
The composition and properties of alumina and ZrO2 based biomaterials are listed in Table 1:
PROPERTY UNITS ALUMINA MG-PSZ TZP
Chemical
composition
99.9% Al2O3 +
MgO
ZrO2 + 8-10
mol % MgO
ZrO2 + 3 mol
% Y2O3
Density Gcm-3 ≥3.97 5.74-6 >6
Porosity % <0.1 - <0.1
Bending
strength
MPa >500 450-700 900-1200
Compression
strength
MPa 4100 2000 2000
Young modulus GPa 380 200 210
Fracture
toughness Kic
MPa m-1 4 7-15 7-10
Thermal
Expansion
Coeff.
K-1 8 x 10-6 7-10 x 10-6 11 x 10-6
Thermal
conductivity
W mK-1 30 2 2
Hardness HV 0.1 2200 1200 1200
Table 1 (given from Piconi, 1999 )
Table 1 shows that zirconia ceramic exhibits higher bending strength (Wagner and Chu,
1996; McLaren, 1997) and fracture toughness (Wagner and Chu, 1996) than alumina
ceramics. Additionally, its Young modulus is much lower than that of alumina, in the same
order of magnitude of stainless steel alloys (CoCr alloy 230 GPa), pointing out its interesting
elastic deformation capability. Fracture toughness is a very important physical property since
it represents the ability of a material to resist crack growth. Clinically, lots of subcritical loads
are applied on the materials by chewing, leading to the growth of subcritical cracks.
Therefore, materials with higher fracture toughness are more ideal clinically, since it takes
more energy to cause crack growth (McLaren and Terry, 2002).
15
LITERATURE REVIEW
2.1.4.3 Zirconia products
Partially Stabilized Zirconia (PSZ) is a product consisting of pure zirconia and stabilizing
oxides like CaO, MgO, CeO2, Y2O3. Its microstructure at room temperature consists of cubic
zirconia as the major phase, with monoclinic and tetragonal zirconia precipitates as the minor
phase (Subbarao, 1981). It has been observed that tetragonal metastable precipitates, finely
dispersed within the cubic matrix, were able to transform into the monoclinic phase when the
constraint exerted on them by the matrix was relieved (i.e. by a crack advancing in the
material). In that case, the stress field associated with expansion due to the phase
transformation acts in opposition to the stress field that promotes the propagation of the crack.
An enhancement in toughness is obtained, because the energy associated with the crack
propagation is dissipated both in the T-M transformation and in the process of overcoming the
compression stress due to the volume expansion (Garvie, 1972; Garvie, 1975). Several PSZs,
like Y2O3-ZrO2 or MgO-ZrO2, were tested as ceramic biomaterials. Mg-PSZ (8% mol MgO in
ZrO2) showed favorable results, but its application diminished rapidly due to the rather coarse
grain size (in the range 30-40µm), the resultant high residual porosity, and the higher sintering
temperatures compared to that for TZP materials (Tetragonal Zirconia Polycrystals).
Additionally, difficulties in obtaining Mg-PSZ precursors free of SiO2, Al2O3 and other
impurities (Leach, 1987) together with the increase in SiO2 contents due to the wear of the
milling media during powder processing before firing (Rühle, 1984), have contributed to the
shift in interest towards TZP materials. Ceramics containing MgO and magnesia silicates,
such as MgSiO3 and Mg2SiO4, may form at the grain boundaries, lowering the MgO contents
in the grains and promoting the formation of the monoclinic phase, which in turn leads to a
further reduction of the mechanical properties and stability of the material in a wet
environment (Leach, 1987).
Tetragonal Zirconia Polycrystals (TZP) ceramic is composed mostly out of the T-phase at
room temperature and contains approximately 2-3% Y2O3 as stabilizing factor. The fraction
of T-phase retained at room temperature is dependent on the size of the grains, on the yttria
content and on the grade of constraint exerted on them by the matrix (Rieth, 1976; Gupta,
1978). The tetragonal grains show a metastable nature. A critical grain size exists, linked to
the yttria concentration, above which spontaneous T-M transformation of grains takes place;
whereas this transformation would be inhibited in a grain structure that is too fine
16
LITERATURE REVIEW
(Theunissen, 1992). Surface tetragonal grains are not constrained by the matrix, and can
transform to monoclinic spontaneously or because of abrasive processes that induce
compressive stresses at a depth of several microns under the surface (Reed, 1977).
Aluminosilicate glasses in the grain boundaries scavenge yttrium ions from TZP grains,
leading to a loss of stability of the tetragonal phase (Lin, 1990). Moreover, mullite (3Al2O3
2SiO2) pockets were detected in the aluminosilicate glass, which lead to a loss of material
stability in a wet environment.
2.1.4.4 Aging of zirconia ceramics
The mechanical property degradation in zirconia, known as ‘aging’, is due to the progressive
spontaneous transformation of the metastable tetragonal phase into the monoclinic phase. This
behavior is well known at a temperature range above 200°C and in the presence of water or
vapor (Sato, 1985a, b).
The aging steps of TZP as given by (Swab, 1991) are:
1. The most critical temperature range is 200-300°C.
2. Aging reduces strength, toughness and density of the material, and increases the
monoclinic phase content.
3. Degradation of mechanical properties is due to the T-M transition, which takes place
with micro and macro cracking of the material.
4. T-M transition starts on the surface and progresses into the bulk of the material.
5. Reduction in grain size and/or increase in concentration of stabilizing oxide reduce the
transformation rate.
6. T-M transformation is enhanced in water or in vapor.
The variability in aging behavior among different zirconia materials is related to the
differences in equilibrium of the microstructural parameters, such as concentration and
distribution of yttria grain size, and population and distribution of flaws (Lilley, 1990). Stable
performances of TZP ceramics in a wet environment were reported by several authors
(Chevalier, 1977; Swab, 1991; Shimizu et al, 1993; Burger, 1995; Fujisawa, 1996; Burger,
1997; Geis-Gerstorfer and Fässler, 1999). Hence, there is experimental evidence that TZP
stability can be controlled acting on several parameters, such as stabilizing oxide
concentration, distribution, grain size and residual stresses in the ceramics (Lepistö, 1992), or
the presence of the cubic phase (Chevalier et al, 2004).
17
LITERATURE REVIEW
The degradation resulting from aging is characterized by surface roughening and
microcracking at the surface (Chevalier, 2006). Garvie (1975) first pointed out that grinding
increases the strength of ceramics containing metastable tetragonal zirconia compared with
fine polishing. Another recent study showed that the grinding of 3Y-TZP ceramics induced no
monoclinic phase formation, but only a rhombohedral zirconia and a strained tetragonal
zirconia phase formation (Denry and Holloway, 2006). This led to a significant increase in
mean flexural strength and increased resistance to crack propagation, but was also associated
with surface and subsurface damage, with formation of microcraters and grain pullout.
Although annealing successfully reversed the zirconia transformation, the surface and
subsurface damage created by grinding remains and could lead to failure by crack propagation
(Denry and Holloway, 2006). Similarly, another group of researchers tested the influence of
surface and heat treatments on the flexural strength of Y-TZP ceramics (Guazzato et al,
2005b) and In Ceram Zirconia® (Guazzato et al, 2005a). In both studies it was concluded that
sandblasting and wet grinding did increase the flexural strength of the ceramics, due to the
monoclinic transformation, but also led to microcracking and strength degradation. Hence, it
was suggested that any surface treatment performed on In-Ceram Zirconia should always be
followed by heat treatment to avoid strength degradation (Guazzato et al, 2005a), while in the
case of Y-TZP ceramics, an initially weaker (with no surface treatment) but in the long-term
more stable (no strength degradation) material may be more desirable (Guazzato et al,
2005b).
The aging sensitivity of Y-TZP is directly linked to the type (compressive or tensile) and
amount of residual stresses. Rough polishing produces a compressive surface stress layer
beneficial for the aging resistance, while smooth polishing produces preferential
transformation nucleation around scratches, due to elastic/plastic damage and the tensile
residual stresses occurred (Deville et al, 2006).
Another relevant aspect for the stability of the material in a biological environment is the
presence of glassy phases formed by SiO2, Al2O3, TiO2 or CaO impurities in grain boundaries.
These impurities may come from the chemical precursors, from the milling bodies used in
powder processing, or may be added to powders as sintering aids. Their presence leads to a
loss of stability of the tetragonal phase, as it was demonstrated that aluminosilicate glassy
phases in grain boundaries are able to scavenge yttrium ions from TZP grains (Lin, 1990).
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LITERATURE REVIEW
Biological safety of zirconia:
An in vitro study reported that Y-PSZ shows a dose dependent cytotoxicity; its toxic effect is
similar to that of alumina, and both lower than that of TiO2 (Dion, 1994). In-vivo studies have
shown an absence of local or systemic toxic effects after the implantation of zirconia ceramics
into muscles or bones of different animals or after powder injection in mice (Bukat, 1990;
Richter, 1994; Walter, 1994). During tests, especially in the early postoperative phase,
connective tissue is frequently observed at the bone-ceramic interface (Tateishi, 1994).
2.2 Classification of high-strength all-ceramic systems:
High-strength ceramic core materials may be classified according to their chemical structure
into 3 major groups (Raigrodski, 2005):
2.2.1 Glass-ceramics
They are multiphase materials that contain an amorphous, glassy phase and crystalline
constituents.
2.2.1.1 Leucite Reinforced Glass-Ceramics
The main representatives of this category are the IPS Empress (Ivoclar Vivadent®, FL-
Schaan) and the Optec® OPC (Jeneric Pentron, D-Kusterdingen). These core materials use
crystalline filler to reinforce glass-ceramic structures. Copings may be fabricated by using
either a heat-pressing procedure or via CAD/CAM technology. The restorations are highly
translucent (Heffernan et al, 2002b, a) providing the potential for a highly esthetic restoration.
Therefore, they are not recommended for cases where the underlying abutment is a discolored
tooth, a metallic-core built up, or a metal implant abutment. The reported flexural strength of
this core material ranges between 105-120 MPa, and the fracture toughness from 1.5 to 1.7
MPa x m1/2 (Campbell, 1989; Seghi et al, 1990; Seghi et al, 1995; Seghi and Sorensen, 1995).
The strength of these restorations depends on a successful bond to the tooth structure and,
therefore must be adhesively cemented. Their indication is restricted only for veneers or
crowns at the front region giving survival rates up to 95% after 11 years of clinical service
(Fradeani and Redemagni, 2002).
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LITERATURE REVIEW
2.2.1.2 Lithium Disilicate Glass-Ceramics
The main representative of this category is the Empress II® (Ivoclar, Schaan, Liechtenstein)
core material. The framework can be fabricated either with the lost-wax and heat-pressure
technique, or can be milled out of prefabricated blanks. Its flexural strength ranges from 300
to 400 MPa (Schweiger, 1999) and its fracture toughness between 2.8 and 3.5 MPa/
m1/2(Schweiger, 1999; Quinn et al, 2003). It is recommended that these restorations should be
etched and adhesively luted to enhance their strength and longevity (Sorensen, 1999). The
material is indicated not only for the fabrication of anterior FPDs, but also for short-span
posterior FPDs (pontic not wider than a premolar) extending up to the second premolar
(Sorensen, 1999; Holand et al, 2000). Esquivel-Upshaw et al (2004) reported a survival rate
of 93% for posterior Empress II FPDs after 2 years. Marquardt and Strub (2006) reported a
survival rate of 100% for single crowns and 70% for FPDs extending up to the second
premolar after 5 years of function.
2.2.2 Glass-infiltrated ceramics
These products consist of infiltrating molten glass to partially sintered oxides. The main
representatives of this category are In-Ceram Alumina®, In-Ceram Spinell® and In-Ceram
Zirconia® (Vita, D-Bad Säckingen).
2.2.2.1 In-Ceram Alumina® (Vita, D-Bad Säckingen)
The material is composed of a highly sintered-alumina glass-infiltrated core and the veneering
porcelain. The fabrication of the core/framework can be carried out either with the slip-cast
technique or by the milling out of prefabricated partially sintered blanks through CAD-CAM
technology. The flexural strength of the material ranges between 236 and 600 MPa (Giordano
et al, 1995; Guazzato et al, 2002) and the fracture toughness between 3.1 and 4.61 MPA/m1/2
(Seghi et al, 1995; Wagner and Chu, 1996). It is recommended for anterior and posterior
crowns, as well as for 3-unit anterior FPDs (Sorensen, 1992; McLaren, 1998). Because of its
semiopaque core, the ceramic does not allow full transmission of light and provide therefore
limited esthetic results (Heffernan et al, 2002b, a).
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LITERATURE REVIEW
2.2.2.2 In-Ceram Spinell® (Vita, D-Bad Säckingen)
The In-Ceram Spinell® consists of a MgAl2O4 core infiltrated with glass. The fabrication
procedures are the same as those for In-Ceram Alumina. Its flexural strength is lower than
that of In Ceram Alumina ranging between 283 and 377 MPa (Magne and Belser, 1997;
McLaren, 1998; Schweiger, 1999), but its translucency is twice as high. Therefore, it is
indicated for anterior crowns, where esthetic demands are higher (Fradeani and Redemagni,
2002).
2.2.2.3 In-Ceram Zirconia® (Vita, D-Bad Säckingen)
The In-Ceram Zirconia® core consists of glass-infiltrated alumina with 35% partially
stabilized zirconia. Its flexural strength ranges from 421 to 800 MPa and its fracture
toughness from 6 to 8 MPa x m1/2 (McLaren, 2000; Chong et al, 2002; Guazzato et al, 2002).
The fabrication may be carried out either with the slip-casting technique or with CAD/CAM
technology. The high opacity of its core (Heffernan et al, 2002b, a) restricts its application
only for the fabrication of posterior FPD’s, resulting in successful short-term data (Suarez et
al, 2004).
2.2.3 Polycrystalline ceramics
This category contains materials with densely packed particles and no glassy components.
They cannot be processed into shapes without the use of Computer-Assisted-
Design/Computer-Assisted-Machining (CAD/CAM) technologies.
2.2.3.1 (CAD-)/CAM Systems:
2.2.3.1.1 Definition/Historical Background
The term CAD/CAM, which comes from machine-tool technology and stands for “Computer-
Aided-Design / Computer-Aided-Manufacturing”, designates the three-dimensional planning
of a workpiece on the screen of a computer with subsequent automated production by a
computer controlled machine tool (Tinschert et al, 2004a). In 1971, Francois Duret
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LITERATURE REVIEW
introduced CAD-CAM technology to the field of dentistry (Duret et al, 1988). His idea was
based upon the assumption that the technologies established in industry could be easily
transferred to dentistry. The industrial use of CAD-CAM allows the production of any
number of similar workpieces automatically, while saving time and manual effort. In dental
medicine, however, this philosophy can not be applied due to the demands of the individual
adaptation of the restoration design (one-of-a-kind production) to the patient (Tinschert et al,
2004a).
2.2.3.1.2 CAD-CAM Components
The contemporary CAD/CAM systems consist of three components (Luthardt, 2001a, b):
1. The scanner, which scans the dental preparation provided by the dentist either intraorally or
extraorally by reference to tooth models. For inlays and single crown frameworks, just the
surface data of the prepared teeth need to be digitized. For FPD frameworks or additional
occlusal characterization, further data from the neighboring teeth and antagonists, as well as
from the spatial relation of the prepared teeth to one another, are required.
2. The software CAD consists of a computer unit used for the three-dimensional planning and
design of restorations on the computer screen. The software programs available today offer a
high level of intervention and permit the design and production of an individually adapted
restoration.
Systems not offering a full CAD component are not considered as CAD/CAM systems but
just as CAM systems. Therefore, we can refer to them as (CAD-)/CAM systems (Witkowski,
2005).
3. The hardware CAM covers different production technologies for converting the virtual
restoration into a dental material. At present, computer-controlled milling or grinding
machines are mainly used. They machine the restoration from the full material block
consisting of prefabricated metal or ceramic. As a rule, after the CAM production, some
manual corrections and final polishing or individualization of the restoration with staining
colors or veneering materials are required to be carried out by the dental technician (Luthardt,
2001a, b).
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LITERATURE REVIEW
2.2.3.1.3 Open/Closed (CAD-)/CAM systems
Most CAD-CAM systems in dental technology operate as closed data systems, i.e., all
components, such as the scanner, the CAD and CAM units, are linked by the specific data
format of the user. The materials used for producing the restorations are also part of this
compound, in the sense that code systems are used.
On the other hand, more and more CAD-CAM systems operating with an open data exchange
are being introduced in the dental market. In this case, the 3-D volume model of the design is
transferred from CAD to CAM in a neutral data format. This language is an industrially
compatible format (such as stereolithography language [STL]), which allows free choice
among different production centers and CAM systems (Witkowski, 2005).
2.2.3.1.4 Materials
The material groups available for the various CAD-CAM systems are as follows:
Silicate ceramics; glass-infiltrated aluminium oxide ceramics; densely sintered aluminium
oxide ceramics; densely sintered zirconium dioxide ceramics ( ZrO2 Y-TZP Zirconia, Yttria-
Tetragonal-Zirconia-Polycrystal), manufactured as green stage, presintered stage and
completely sintered stage; titanium; precious alloys; nonprecious alloys; acrylics of improved
strength and castable acrylics (Witkowski, 2005).
The Procera AllCeram® (Nobel Biocare, S-Göteborg) is a polycrystalline ceramic consisting
of a densely sintered high-purity aluminium-oxide core (Oden et al, 1998). It has a flexural
strength between 500 and 650 MPa (White et al, 1996; Zeng et al, 1996)and a fracture
toughness of 4.48-6 MPa x m1/2 (Christel et al, 1989; Wagner and Chu, 1996). It is
recommended for the fabrication of anterior and posterior crowns, but its use for 3-unit FPDs
is still questionable (Raigrodski, 2005).
Zirconium dioxide has been introduced into dentistry as a framework material for various
indications. The ZrO2 frameworks for crowns and FPDs are made by milling in the green
stage (diamonds with cooling liquid) (Filser, 1997), the presintered stage(dry carbide burs),
and the completely sintered stage (diamonds with cooling liquid) (Witkowski, 2005). ZrO2 that
belongs to the green stage group can be individualized by coloring of the framework
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LITERATURE REVIEW
according to the Vita shade concept. Erdelt (2004) showed no changes in the physical
properties of the materials when colored by an oxide liquid prior to the sintering process.
2.2.3.1.5 Yttrium Tetragonal Zirconia Polycrystals (Y-TZP)
Y-TZP is a glass-free, high-strength polycrystalline ceramic material with a flexural strength
of 900 to 1200 MPa and fracture toughness of 9 to 10 MPa x m1/2 (Christel et al, 1989). It is
indicated for anterior and posterior crown copings and FPD frameworks (Luthardt et al,
2004). The majority of the Y-TZP –based (CAD-)/CAM systems use CAM of partially
sintered Y-TZP blanks (Lava, 3M Espe Dental AG, Seefeld; Cercon, DeguDent, Hanau;
Cerec InLab, Sirona Dental Systems, Bensheim; Procera AllZirkon, Nobel Biocare, S-
Göteborg). The size of partially-sintered infrastructures is increased during the milling stage
to compensate for prospective shrinkage (20-25%) occurring during final sintering
(Raigrodski, 2005). The milling of these blanks is faster and results in less wear and tear to
the hardware (Raigrodski, 2004b). With fully sintered blanks, such as DC-Zirkon (DCS-
Precident, DCS Dental AG, CH-Allschwill), there is no shrinkage involved in the milling
process, but microcracks may be introduced to the infrastructure (Luthardt et al, 2004).
Product examples of ZrO2 materials and the groups according to the milling/grinding
technology are: (Witkowski, 2005)
Milling at green stage: Cercon Base (Cercon), Lava Frame (Lava), Hint-Els Zirkon TZP-G
(DigiDent), ZirkonZahn (Steger), Xavex G 100 Zirkon (etkon)
Grinding at presintered stage: In Ceram YZ-Cubes (Cerec InLab), ZS-Blanks (Everest),
Hint-Els Zirkon TZP-W (DigiDent), DC-Shrink (Precident DCS)
Grinding at completely sintered stage: DC-Zirkon (Precident DCS), Z-Blanks (Everest),
Zirkon TM, Pro 50 (Cynovad), Hint-Els Zirkon TZP-HIP (DigiDent), HIP Zirkon (etcon)
Only a few CAD-CAM systems offer the possibilities of using different materials and
fabricating occlusal surfaces. Even if a complete reconstruction of the occlusal surface
(framework production only) is not wanted, a framework design according to anatomical
aspects with inclusion of the contact relations should be a primary goal. In this so called
“intelligent framework design”, the construction is strengthened in all areas with sufficient
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LITERATURE REVIEW
clearance from the antagonists, the neighboring teeth and the gingiva, so that the veneering
ceramic can be fired on with uniform thickness (Rudolf, 2003). This procedure should ensure
that the veneering ceramic receives sufficient support while avoiding the occurrence of too
thick veneer layers and material stresses because of layer thickness fluctuations. In employing
this method, the risk of veneer spalling off is also reduced (Tinschert et al, 2004a).
2.2.3.1.6 The Cerec System
In 1985, the Cerec I (Brains, Zürich, Switzerland) CAD-CAM system was introduced to the
dental market. In 1994, Siemens (Bensheim, Germany) introduced the Cerec 2 unit. Due to
the restricted efficiency of the computer at that time, the full effects of the “correlation” and
“function” construction modes were limited (Mörmann et al, 1999). The Cerec 3 system
(Sirona, Bensheim, Germany) was introduced in 2000. After 1 year of use, hardware and
software improvements were implemented in early 2001. The chairside Cerec 3D system is an
improved version of the Cerec 2; including the intraoral 3D scanning camera, image
processing, computing power and a form-grinding unit. With this advance in computer
efficiency, the two-impression “correlation” and “function” modes for designing partial and
full crowns are able to proceed as desired, using occlusion and preparation optical
impressions without loss of time (Mörmann and Bindl, 2002). Consequently, the occlusion
and the preparation images can be used alternately to fit design suggestions arising from the
morphologic data bank to the individual situation. The separate form-grinding unit, working
true to morphologic detail and with fine surface quality, is connected to the optical unit by
radio control. The form-grinding unit receives data from the control unit, independent of its
location in the office. The next restoration can be designed while the first is being milled
(Mörmann and Bindl, 2002). The form-grinding unit is fitted with a laser scanner (Cerec
Scan, CerecinLab Sirona) and can be used by itself with a standard personal computer for
indirect application in the dental laboratory. In April 2001, the application was expanded for
the fabrication of three-unit fixed partial denture frames (Mörmann and Bindl, 2002).
2.2.3.1.7 CAM Technologies
The CAM technologies can be divided in three groups according to the technique used
(Witkowski, 2005):
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1. Subtractive Technique from a Solid Block
The CAM technique most commonly applied in manufacturing frameworks for single
crowns and FPDs is to cut the contour out of an industrially prefabricated , solid block of
different materials (Andersson et al, 1989; Witkowski, 2002). The size of the material
blocks available for the milling units limits the size of the FPD. When industrial
prefabricated zirconium dioxide blocks are used, the restoration can be shaped both before
and after the block is sintered. Consequently, we can have the green machining process of
presintered ceramic blocks and the hard machining process of densely sintered ceramic
blocks. In regard to the green machining, it offers the benefit of saving time and grinding
tools for the labor, but the sintering shrinkage that occurs is difficult to be computer-
controlled for extensive restorations. Further, it has not been proven whether or not the
grinding dust arising in green machining leads to damage of the milling unit in the long
run. The hard machining on the other hand is time-consuming, leads to greater wear of the
grinding tools, and there is a risk of introducing unwanted surface or structural defects
into the ceramic during the machining (Tinschert et al, 2004a).
The DCS Precident system (Allscwill, Switzerland) is based on the hard machining
process, using a laser scanner to scan multiple units at once, and software which suggests
connector sizes and pontics for frameworks. The system uses a variety of materials
including porcelain, glass-ceramic, In Ceram, densely sintered Zirconia (DC-Zirkon),
metals and fiber reinforced composite. There’s no shrinkage or sintering involved after
milling (Giordano, 2003).
The Cercon system (Degudent, Hanau, Germany) is not a CAD/CAM system. It
requires a wax-up of the desired bridge framework. This wax-up is then scanned and
through software manipulation and CAM processing an oversized coping of partially
sintered Zirconia is milled out. This oversized coping (compensation for the 25-30%
sintering shrinkage) will afterwards be fired for 6-8 hours at high temperature in order to
produce a fully sintered zirconia.
The Lava system (3M/ESPE Dental AG, Seefeld, Germany) is an offsite system. The
central unit uses an optical scanner to scan multiple units at once. The software
automatically finds the margin, suggests pontics and designs the desired framework
(Giordano, 2003). Afterwards, with the milling machine, an oversized coping from
partially sintered zirconia is milled out to compensate for sintering shrinkage. An
additional feature is the ability to color the zirconia by dipping it in various solutions prior
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LITERATURE REVIEW
to dense sintering. The entire procedure from scanning to milling is completed at the
center and then returned to the labor (Giordano, 2003).
2. Additive Technique by Applying Material on a Die
There are three different systems that apply the framework material on a die of a prepared
tooth (Witkowski, 2005):
Procera (Nobel-Biocare AB, Göteborg, Sweden)
The first system that was based on the knowledge of exact dimensional changes that take
place during sintering was the Procera system (Nobel Biocare), which Andersson and
Oden introduced in 1993 (Andersson and Oden, 1993). The system was also the first to
introduce an industrialized process in which the framework is manufactured in a remote
production unit (Anusavice, 1989; Andersson et al, 1998). The scanner in the dental
laboratory scans the working die, and stores the information in a computer. After
scanning, the technician marks the preparation margins on the computer screen and
indicates the desired material (alumina or zirconia) framework thickness, and, in some
instances, different opacities. This information is then compressed and transferred via a
modem line to the production unit, which uses the information to calculate the anticipated
shrinkage and fabricate an enlarged die. Alumina or zirconia is dry pressed against the
enlarged die, and the temperature is raised to a temperature similar to the presintering
stage. At this point in the process, the enlarged and porous coping is stable. Its outer
surfaces are milled to the desired shape and the coping, removed from the enlarged die,
and sintered into the furnace for firing to full sintering. During this cycle, the coping
shrinks to fit the dimensions of the original working die. The completed coping is then
sent back to the laboratory, where it is veneered with the compatible silica-based ceramic
(Sadan et al, 2005).
The second system (EPC 2019, Wol-Ceram System, Wol-Dent, Ludwigshafen, Germany)
(Wolz, 2002) of this group generates the ceramic powder (In-Ceram Alumina and
Zirconia, Vita Zahnfabrik, Bad Säckingen, Germany) directly on the die of the master
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LITERATURE REVIEW
model. The ceramic material (slurry stage) is generated by an electrophoretic dispersion
method within a few minutes on the die(s) for single copings and FPDs. Overextended
margin contours can be manually trimmed and the outside contour is shaped by a CAM
process. Then, the coping is removed from the die glass-infiltrated at high temperature
(1140°C) (Witkowski, 2005).
The third system involves the solid direct form fabrication technique, which generates
copings and frameworks for FPDs of pure Al2O3 and ZrO2 ceramics in a production center
(ce.novation, Inocermic, Hermsdorf Germany). The dispersed super-fine nanoceramic
powders consist of particles well below 100 nm in diameter. With this technology, the
frameworks attain high strength and calculable sintering shrinkage (Brick, 2003). These
new technologies are relatively new and need further development (Witkowski, 2005).
3. Solid free form fabrication
This category includes new technologies originating from the area of rapid prototyping,
which have been adapted to the needs of dental technology (Gebhardt, 2000; Wohler,
2003). The first system applying this technology for dental use was the wax plotter
technique, which works according to the ink jet principle. The machine builds (solid free
form) frameworks and full crowns in wax for the casting technique in alloys and titanium
(Wax Pro 50, Cynovad, Montreal, Canada) (Witkowski, 2005). A second technology
originating from rapid prototyping is the stereolithography (Perfactory, Delta Med,
Frieberg, Germany). In this technique, the restoration is produced from light sensitive
plastic, which can be converted into any desired alloy with the casting technique
(Witkowski, 2003). Occlusal splints and diagnostic templates for oral implantology can
also be produced with this technique. The third technique is the selective laser sintering,
where sinterable powder materials are built up to form 3-D restorations. Every applied
powder layer is fused by means of a laser (Medifacturing, Bego Medical, Bremen,
Germany) (Strietzel, 2001). Only metals can be processed currently, whereas laser
sintering of ceramics is still in the testing phase.
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2.2.3.1.8 Marginal fit of CAD-CAM restorations
The fitting accuracy of a CAD-CAM produced restoration depends primarily upon the
quality of the digital data acquisition and processing, the filters matched to the measuring
system and used for removing measuring-induced noise and stray points; and upon the
position of the preparation border, equator and undercuts. Therefore, the number of
measuring points is not an absolute criterion for the quality of the data record (Rudolf,
2003). However, precise tooth preparation and careful design are still prerequisites for
good fitting accuracy of a CAD-CAM produced restoration (Luthardt, 1998; May et al,
1998; Hertlein, 2001; Tinschert et al, 2001b; Beuer, 2003; Nakamura et al, 2003).
2.3 Clinical and technical aspects of all-ceramic FPDs
2.3.1 Clinical aspects
2.3.1.1 Preparation design
Although there is no standard, the preparation design for all-ceramic FPDs requires an
occlusal reduction of 2 mm (Banks, 1990) and a deep chamfer or circular shoulder with an
inner rounded angle between 1.0 and 1.5 mm wide. The occlusal angle of convergence should
not be greater than 10 degrees (Friedlander et al, 1990; Sorensen et al, 1998; McLaren and
White, 1999, 2000).
While 1.4 to 1.7 mm of facial reduction is necessary to achieve an esthetic result with metal-
ceramic systems (Sorensen, 1999; Sorensen et al, 1999), a reduction of 1.0 to 1.5 mm is
required for ceramic systems (Chiche, 1994; Rosenstiel, 1995; Shillingburg, 1997).
2.3.1.2 Translucency/Esthetics
All ceramic restorations offer improved esthetic results because of their natural reflection and
translucency of light. A recent study (Heffernan et al, 2002a, b) compared the translucency of
6 all ceramic system core materials at clinically appropriate thicknesses. In order of
decreasing translucency, the ranges were as follows: Vitadur Alpha Dentin(standard)>In-
Ceram Spinell>Empress, Procera, Empress II>In-Ceram Alumina>In-Ceram Zirconia, 52 SF
alloy. The authors demonstrated that all systems showed an increase in translucency after the
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veneering process and glazing cycle, with the exception of In-Ceram Zirconia and the metal-
ceramic specimens, which showed no improved opacity after the glazing cycle. Another study
evaluated the light transmission through different types of ceramic frameworks with different
cements (zinc phosphate/adhesive cement). The light transmission coefficients of the
infrastructures (0.9 mm) are listed as follows: In-Ceram Spinell>Empress II>Procera>Y-
TZP>In-Ceram Alumina>In-Ceram Zirconia (Edelhoff, 2002). When the thickness of a
material increases from 0.5 to 1.5 mm, the opacity rises from 65 to 85% (Hauptmann, 2000).
The knowledge of the opacity limitations of each system is of great importance, since it
allows us to successfully determine, especially for the anterior region, which system achieves
better esthetics.
2.3.1.3 Fracture resistance testing
The physical properties of newly developed dental ceramics must be tested before they can be
recommended for clinical application. Generally, a study design which uses fixed abutments
produces fracture loads that are much higher than those reported in studies which employ
mobile abutments (Kappert et al, 1991; Kelly et al, 1995). The fracture loads thus obtained
should be viewed critically in terms of their clinical relevance.
In vitro investigations, finite element analyses (FEA) of FPDs and fractographic studies
conducted on clinically failed FPDs have shown that the stress distribution, fracture origin
and mode of failure of all-ceramic FPDs are substantially different from those of crowns
(White et al, 1994; Kelly et al, 1995; Zeng et al, 1998; Proos et al, 2000; Thompson, 2000;
Wakabayashi and Anusavice, 2000; Chong et al, 2002). In the case of crowns, tensile stresses
are distributed at the inner surface, while all-ceramic FPDs tend to fail at the connector area
where there is a peak of tensile stress. Consequently, the properties of the porcelain which
veneers the stronger core material control the failure fate of the connector (Kelly et al, 1995;
Proos et al, 2000). Therefore, some authors concluded that all areas of the restoration that are
subjected to tensile stress should not be veneered with porcelain (White et al, 1994; Zeng et
al, 1998). Another study showed that increasing the connector height from 3 to 4 mm
dramatically reduced the stress levels within the connectors (Kamposiora et al, 1996). More
recent studies carried out by Guazzato et al (2004a, b, c, d) confirm the previous findings
indicating that the material subjected to the peak of tensile stress dictates the ultimate strength
of the restoration. Some fractures, observed at the interface between the core ceramic and the
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LITERATURE REVIEW
veneering porcelain, have been related to the stress enhancement arising from large
differences in elastic modulus between the veneer and the core ceramic materials.
Contact loading, a new approach in the testing of dental ceramics, has been suggested by
Lawn et al (2001) to evaluate the physical properties of ceramics (Peterson et al, 1998a;
Peterson et al, 1998b; Jung et al, 1999a, b; Kim, 1999; Jung et al, 2000).
A summary of the contact loading research studies has indicated the following:
1. Contact damage mode analysis for fine-grain, low-toughness dental porcelain found
that the fracture mode was conventional hertzian cone fracture. Fractures in high
crystalline, higher toughness, coarse-grain ceramics were caused by quasi-plastic
deformation.
2. Stress-indentation curves of hertzian contact damage confirmed the ranking from
flexural strength data (i.e., micaceous glass-ceramic<porcelain<alumina<zirconia).
3. For cyclic contact tests in dry versus wet environments, and for any given contact
load, the strength of a ceramic degraded with an increase in the number of loading
cycles. Most ceramic materials degraded significantly between 10.000 to 100.000
cycles. The presence of water enhanced damage accumulation in cyclic indentation.
4. Machining effects were shown to cause surface damage and reduce surface strength
but to have only a secondary effect on the initiation and evolution of cone cracking
and quasi-plasticity.
In regard to the mode of fracture and its origin, new generations of multilayered core-
veneer ceramics are of increased interest. Jung et al (1999b) reported that the substrate has
a profound influence on the damage evolution, which leads to ultimate failure in the
bilayered systems. Thompson (2000) showed that the different failure modes and failure
origins by ceramic testing are dependent on testing methodology, relative layer heights,
and very likely, the ceramic system being investigated. White et al (1994) demonstrated
that layered ceramics made of strong cores veneered with weaker feldspathic porcelain
may be prone to failure when the feldspathic porcelain surfaces are subjected to tensile
force. Wakabayashi and Anusavice (2000) found that as the elastic modulus of the metal
substrate increased, there was an increase in the load to cause failure, but no change in
fracture origin. As the thickness ratio of ceramic core to veneer increased, the site of crack
initiation shifted from the veneer porcelain to the core porcelain. The mean fracture
strength increased as the core-to-veneer thickness ratio increased, but it did not exceed
that of the core material.
31
LITERATURE REVIEW
2.3.1.4 Marginal Fit
The accuracy of fit of all-ceramic restorations is important for the integrity of dental and
periodontal tissues, dissolution of luting agents and the fracture resistance of the restorations.
Poorly fitted dental restorations are believed to be closely associated with the development of
secondary caries and periodontitis (Gardner, 1982; Lang et al, 1983; Knoernschild and
Campbell, 2000; Becker and Kaldahl, 2005). Up to this date, the dental research has
demonstrated a wide variety of values for clinically acceptable marginal gaps of dental
restorations. These gaps vary from 20 to 200 µm (Dreyer, 1958; McLean and von Fraunhofer,
1971; Rehberg, 1971; Marxkors, 1980; Fransson et al, 1985; Spiekermann, 1985;
Kerschbaum, 1995). In 1985 Spiekermann estimated a marginal gap up to 100 µm as
clinically acceptable. In 1995, Kerschbaum supported a marginal gap value of up to 200 µm
as clinically acceptable for dental restorations.
In vitro results demonstrating marginal gaps of 64-83 µm in CAD/CAM-generated all-
ceramic restorations are quite promising (Sulaiman et al, 1997; Lin et al, 1998; Bindl et al,
1999; Boening et al, 2000; Tinschert et al, 2001b). An in vivo study by Procera titanium
crowns yielded gap widths that were 61-70 µm wider in the bucco-lingual direction and 58-73
µm wider in proximal locations than gap widths measured in vitro (Karlsson, 1993). Mean
values between 64-74 µm were reported for zirconia multi-unit frameworks produced by the
DCS CAD/CAM system (DCS, Allschwil, Switzerland) (Tinschert et al, 2001b). Another
study reported mean marginal gaps of 75 µm for the DigiDent (Girrbach, Pforzheim,
Germany) CAD/CAM system (In Ceram Zirkonia blanks) and 65 µm for both the Lava (3M
ESPE, Seefeld, Germany) system (yttrium-stabilized Zirkonia blanks) and the Cerec InLab
(Sirona, Bensheim, Germany) (In Ceram Zirkonia blanks) system (Reich et al, 2005). The
authors concluded that the accuracy of CAD/CAM-generated three-unit FPDs is satisfactory
for clinical use.
2.3.1.5 Occlusal forces
The normal physiological chewing forces on posterior teeth range between 2-150 N (Eichner,
1963; Bates et al, 1976; De Boever et al, 1978; Jäger, 1989). Maximum biting forces that
may occur in the posterior area vary between 300 and 880 N (Bates et al, 1976; Gibbs et al,
32
LITERATURE REVIEW
1986; Kiliaridis et al, 1993). Males show higher bite values than females. For a male, the
maximal bite force is 382 N for the molar region and 108 N for the incisor region, while for
females the values are 216 N and 108 N, respectively (Helkimo et al, 1977).
In several simulating clinical conditions, a load of 49 N is set for the chewing simulator to
simulate the physiological biting force (Krejci et al, 1990; Behr et al, 1999; Kern et al, 1999;
Koutayas et al, 2000).
Further literature data indicate that the maximum biting forces given for natural teeth and
short-span FPDs in the posterior region range between 50 N and 400 N (Helkimo et al, 1977;
Körber, 1983; Hagberg, 1986). In case of bruxism, the biting forces can rise up to 500 N or
880 N (Kelly, 1995, 1997; Kikuchi et al, 1997). Hence, a mean value of 500 N is considered
as the minimum-demanded fracture resistance of a material for application at the posterior
region. Dental ceramics, due to their subcritical crack growth, exhibit 50% of their initial
fracture resistance after fatigue loading. This degradation should be taken into consideration
in determining the minimum demanded fracture resistance and further ensuring the long-term
clinical stability of the restorations (Sorensen et al, 1998; Geis-Gerstorfer and Fässler, 1999;
Marx et al, 2001). Hence, it appears reasonable that an initial fracture resistance of 1000 N
should be demanded for posterior FPDs.
2.3.1.6 Cementation
The surface treatments indicated for glass-infiltrated ceramics are either airborne particle
abrasion with Al2O3 (50-100 µm at 2.5 bar) with the use of a phosphate-modified resin
cement (Panavia 21, Kuraray, Osaka, Japan) (Isidor et al, 1995; Kern and Thompson, 1995;
Kern and Wegner, 1998; Madani et al, 2000), or tribochemical surface treatment (Rocatec
System, 3M ESPE, Seefeld. Germany) in combination with conventional Bis-GMA resin
cement (Sadoun and Asmussen, 1994; Kern and Thompson, 1995; Ozcan et al, 2001). Some
authors recommend Panavia without a silane or bonding agent (Kern and Thompson, 1995),
whereas others suggest a silane coupling agent to increase the saturation properties of the
ceramic substrate (Madani et al, 2000; Ozcan et al, 2001).
The Rocatec System (3M ESPE, Seefeld. Germany) is an effective and user friendly silica-
coating method (Piotrowski, 2001). It includes 2 steps of airborne particle abrasion and the
application of a silane coupling agent (ESPE-Sil; 3M ESPE, Seefeld, Germany) that bonds to
the silica coated surface and to resin. The treatment with the Rocatec Sytem is not effective
33
LITERATURE REVIEW
for glass-phase-free and densely sintered alumina and zirconia ceramics (Kern and Wegner,
1998; Friederich and Kern, 2002).
For densely sintered alumina- and zircionia-based ceramics, such as Procera All Ceram or Y-
PSZ ceramics, a durable resin bond can only be achieved with airborne particle abrasion (50-
110 µm Al2O3 at 2.5-bar) and a phosphate-modified resin cement (Panavia or Panavia 21)
(Kern and Wegner, 1998; Wegner and Kern, 2000; Friederich and Kern, 2002; Hummel and
Kern, 2004). The high-strength aluminium- and zirconium-based core ceramics do not require
adhesive cementation techniques to strengthen the restoration (Sorensen et al, 1998;
Sorensen, 1999; Fradeani, 2000; McLaren, 2000), whereas acid-etching and cementation
with adhesive resins are recommended for silica-based all-ceramic systems (Wohlwend, 1990;
Fradeani and Aquilano, 1997; Sorensen et al, 1998). With Y-TZP-based materials, adhesive
cementation may be used, but is not mandatory; and traditional luting agents, such as glass-
ionomer cements, may be applied (Besimo et al, 2001; Filser et al, 2001; Suttor et al, 2001).
However, some clinical situations, such as compromised retention and short abutment teeth,
require resin bonding and, therefore, adequate ceramic surface conditioning (Burke et al,
2002; Blatz et al, 2003).
While the elimination of postcementation sensitivity remains a clinical objective, all types of
dental cements cause side effects (Sorensen et al, 1998). Johnson recorded a 32% incidence of
immediate postcementation sensitivity for zink phosphate cement and 19% for glass-ionomer
cement (Johnson et al, 1993). A 9.1% incidence of postcementation symptoms by using the
adhesive technology compares favorably to the incidence of symptoms for conventional
cements (Sorensen et al, 1998).
2.3.2 Technical aspects for all-ceramic FPDs
2.3.2.1 Connector dimensions
The maximum strain in fixed partial dentures is systematically located in the connectors
(Kelly et al, 1995). The law of beams is the basis for the proper design of the connectors and
pontic. The deflection of a beam increases as the cube of its length is inversely proportional to
its width, and is inversely proportional to the cube of its height (Dupont, 1968). When the
underlying structure is deformable, as in the case of dental and periodontal structures, the
34
LITERATURE REVIEW
resultant strains bring the anchoring systems together and open the angle of the axis formed
by abutment teeth. The more deformable a structure is, the closer the tensile strength
approaches the cervix. The tensile strength is mainly induced on the cervical side of the
connectors. The 0.15% strain threshold, beyond which rupture is possible, is only reached
when the connector surface is minimal (3.3 mm2). Extending the surface of the connectors,
which is less consistent with periodontal clinical requirements, is not necessary to ensure
resistance to rupture (Augereau et al, 1998). Kamposiora et al (1996) showed that increasing
the connector height from 3 to 4 mm dramatically reduced the stress levels within the
connectors.
Apart from the size of the connectors, their design is also crucial. Generally, connectors with
a round design do not develop stress peaks, and therefore withstand higher pressures.
Rounding off the connector in the mesiodistal axis to provide an adequate radius also prevents
stress peaks, and thereby allows higher pressures to be applied before fracture occurs (Filser,
2003). Dimension requirements of 3x3 mm have been suggested for all-ceramic FPD core-
connectors (Futterknecht, 1990; McLean, 1993). Another suggestion is 4x4 mm (Kappert,
1990) or 3 mm buccolingually and 4 mm occlusogingivally (McLaren, 1998). Furthermore, it
was suggested that a diameter of 4 mm in case of a molar replacement and 3 mm in other
cases should be applied (Vult von Steyern et al, 2001). In order to minimize the risk of
fracture for all-ceramic FPDs clinically, we should always follow manufacturer
recommendations in regard to the minimum critical dimensions demanded for the connectors.
2.3.2.2 Thermal Expansion Coefficient (TEC)
Metals and porcelains should be selected with a slight mismatch in their TEC’s (metal slightly
larger), so that the metal contracts more than the porcelain on cooling from the firing
temperature to room temperature. This mismatch leaves the porcelain in residual compression
and provides additional strength to the restoration (Anusavice, 2003). This philosophy has
been successfully applied for the fabrication of many multilayered all-ceramic restorations.
Considering the fact that the core is put under tensile stress, a relative high flexural strength
and a minimum critical material thickness are required in order to achieve a stable bond
between the core and the veneering material. In the Empress II system, a minimum core
(TEC=10.6x10-6 /K, 350 MPa flexural strength) thickness of 0.8 mm is required, in order to
achieve a reliable bond-strength (10% TEC difference) with its corresponding veneering
35
LITERATURE REVIEW
material (TEC=9.7x 10-6 /K, 100 MPa flexural strength). In-Ceram Alumina shows a slight
TEC-difference ( 5%) between its core (TEC=7.4 x 10-6 /K, 500 MPa flexural strength) and
its veneering material (TEC=7.0 x 10-6 /K, 84 MPa flexural strength), so that a minimum 0.4
mm core thickness is required for a long-term stability (Kappert, 2001). Thermal expansion
coefficients of different Y-TZP framework values given by the manufacturer range from
10x10-6/K (DC-Zirkon®, DCS Precident, Allschwill, Switzerland and Lava®, 3M ESPE,
Germany) to 10.4x10-6/K (Procera Zirconia®, Nobel-Biocare AB, Göteborg, Sweden) and
10.5 x10-6/K (Cercon®, DeguDent, Frankfurt, Germany and Vita In Ceram 2000 YZ Cubes,
Vita Zahnfabrik, D-Bad Säckingen). The corresponding company, for each of these
framework materials, also develops their respective compatible veneering materials.
2.4 Survival rates of all-ceramic FPDs
Clinical Studies
Empress II:
In a short term clinical study, 4 out of 61 Empress II fixed partial dentures (anterior and
posterior region) failed to fracture, showing a failure rate of 6.7% after an observation period
of 1 year (Sorensen et al, 1999). Three of the four restorations that failed had occlusogingival
connector heights that failed to achieve the recommended design standards.
Esquivel-Upshaw et al (2004) showed a 93% success rate for Empress II posterior FPD’s.
One of the two fractures was associated with short connector height (2.9 mm). A further two-
year follow-up for Empress II crowns and FPD’s (replacing up to the 1st premolar) reported a
100% success rate for single crowns, but 50% survival rate for the FPD’s, which tended to
fracture at the connector regions (Taskonak and Sertgoz, 2005). Marquardt and Strub (2006)
reported a 100% survival rate for Empress II crowns and 70% for FPDs (anterior and
premolar region) after an observation period of 5 years. Three out of the six failures did not
follow manufacturer recommendations in regard to the connector dimensions.
In-Ceram Alumina:
In a 3-year clinical study, In-Ceram Alumina FPDs yielded a 0% failure rate for anterior
FPDs, 11% for FPD’s with premolar pontics and 24% for FPD’s with molar pontics. All
FPDs were cemented with glass-ionomer cement. This study confirmed the indication given
36
LITERATURE REVIEW
from the manufacturer recommending the use of In-Ceram Alumina only for anterior FPD’s
(Sorensen et al, 1998). Another clinical study on posterior In-Ceram Alumina FPDs
(replacing premolar and molar teeth) showed a 90% success rate after an observation period
of 5 years, demonstrating that the In-Ceram technique can also be a viable treatment
alternative for the posterior region (Vult von Steyern et al, 2001). In a long-term retrospective
study with anterior and posterior In-Ceram FPDs, the survival rates were 93% and 83% after
5 years and 10 years, respectively (Olsson et al, 2003).
In-Ceram Zirconia:
In a 3-year clinical observation, the survival rate of In-Ceram Zirconia posterior FPDs was
94.5 % (Suarez et al, 2004). Because of the esthetic limitations, resulting from the high
opacity of its core, this material is recommended only for the fabrication of posterior FPDs
(McLaren and White, 1999).
Y-TZP-based restorations:
In a 1 year clinical evaluation, 22 adhesively cemented 3-unit posterior ZrO2 FPDs, milled out
with the DCM-System®, showed a success rate of 100% (Sturzenegger et al, 2000). Another
1-year clinical study reported a success rate of 100% for 20 3-unit and 6 4-unit posterior ZrO2
FPDs, fabricated with the DCS® Precident- System and cemented using zinc-phosphate
(Tinschert et al, 2001c). Twenty-one posterior ZrO2 FPDs, milled out with the Lava® CAD-
CAM system and cemented with glass-ionomer cement, showed a 100% success rate after an
observation period of 8 months (Pospiech et al, 2002). After a 15.5-month mean evaluation
period, 36 posterior and 10 anterior DC-Zirkon FPDs, cemented with zinc-phosphate cement,
showed no framework fractures (100% survival rate) but only one chip-off defect (2.5%) of
the veneering ceramic material (Tinschert, 2002). Fifty-eight 3- to 5-unit ZrO2 FPDs
(replacing premolars and molars), milled out with the Cercon® (Cercon Smart Ceramics,
DeguDent, Germany) system gave a survival rate of 93% after 2 years (Zembic et al, 2002)
and 83% after 3 years of clinical service (Sailer et al, 2003). In these studies, 18% of the
restorations showed no clinically acceptable marginal discrepancies, indicating that a further
refinement of the manufacturing techniques is demanded. Bornemann et al (2003) examined
44 3-unit and 15 4-unit posterior ZrO2 FPDs, milled out with the Cercon® System and
cemented with zinc-phosphate cement. Their results showed no framework fractures but a
3.5% chip-off of the veneering ceramic after 6 months in service. Molin (2003) reported a
survival rate of 67% for 18 posterior and one anterior ZrO2 Denzir® FPDs after an observation
37
LITERATURE REVIEW
period of 2 years. Fifty posterior DC-Zirkon FPDs (33 three-unit, 14 four-unit and 3 five-
unit), milled out with the DCS® Precident- System and cemented with zinc-phosphate cement,
showed no framework fracture, but 6% demonstrated chip-off veneering defects, after 3 years
of clinical service (Tinschert et al, 2004b). In a 1-year clinical observation, 21 conventionally
cemented Cercon® ZrO2 cantilever-FPDs with maximum extension up to 8 mm showed a
survival rate of 100% (Rinke, 2004). In another study, no framework fractures but only a 15%
chip-off veneering defect rate for 20 conventionally cemented 5-unit (minimum 3 retainers)
posterior and anterior DC-Zirkon FPDs were reported after an observation period of 2 years
(Vult von Steyern et al, 2005).
Although the presented short-term clinical results are quite promising, longitudinal clinical
studies are necessary for the assessment of long-term success and for the establishment of
more specific guidelines for the use of zirconia as a material for the fabrication of FPDs
(Raigrodski, 2004b). Considering that metal ceramic FPDs, which are our current standard of
care, demonstrate a cumulative survival rate of 96% after 5 years and 87% after 10 years
(Walton, 2002), all-ceramic FPDs should demonstrate at least a similar survival rate in
clinical studies if they are to be considered as a predictable restorative alternative (Raigrodski,
2004a).
38
AIM OF THE STUDY
3 Aim of the study
The aim of the present in vitro study was:
1. To compare the fracture resistance and mode of failure of posterior zirconia all
ceramic fixed partial dentures, fabricated with 3 different CAD/CAM systems.
2. To evaluate the effect of fatigue loading in the chewing simulator on the fracture
resistance of the tested materials.
39
OUTLINE OF THE STUDY (FIG. 1)
4 Outline of the study (Fig. 1)
Figure 1. shows the outline of the study:
Figure 1
96 teeth (48 premolars + 48 molars)
32 teeth→ 32 teeth→ 32 teeth→ 16 samples 16 samples 16 samples
DCZirkon + Procera + Vita YZ + e max Ceram e max Ceram e max Ceram 16 FPDs 16 FPDs 16 FPDs
8 aged FPDs
8 non-aged FPDs
8 aged FPDs
8 non-aged FPDs
8 aged FPDs
8 non-aged FPDs
Fracture strength test
40
MATERIALS AND METHODS
5 Materials and Methods
5.1 MATERIALS
5.1.1 Abutment Teeth
Forty-eight caries-free human lower premolars and molars, without fractures and/or
hypoplastic defects, were used as abutments. The teeth were obtained directly after extraction
and stored in 0.1% thymol solution throughout the study. The root surfaces were cleaned from
concrements and desmodontal rests ultrasonically and with handscalers.
5.1.2 Materials used for the fabrication of the all-ceramic FPDs
The all-ceramic FPDs were divided into 3 groups of 16 specimens each. Each group was
fabricated with a different framework material. All frameworks of different groups were
veneered with the same ceramic veneering material (e max Ceram®, Ivoclar Vivadent AG,
Schaan, Liechtenstein).
The composition and properties of the materials used for the fabrication of the frameworks of
the FPDs is listed in Table 2.
41
MATERIALS AND METHODS
FRAMEWORK MATERIAL COMPOSITION
(WEIGHT %)
PROPERTIES
Procera Zirconia ® DC-Zirkon® Vita In Ceram
YZ Cubes®
ZrO2(HfO2) approx. 95%,
<5% HfO2
approx. 95% Approx. 95%, <3%
HfO2
Y2O3 4.5-5.4% <5% 5%
Al2O3 +other oxides <0.5% <1% (Na2O) <1% (Si O2)
Al2O3 or other oxides Not provided <0.5 Na2O Not provided
Volumetric weight >6.05 g/cm3 >6.08g/cm3 6.05 g/cm3
Particle size <0.5µm <0.6 µm 300 nm
Vickers Hardness 1200 HV
1200 HV
1200 HV
Melting temperature 2700°C
Not provided 2706°C
Bending strength 1121MPa
900 MPa
>900 MPa
Elasticity modulus 210 GPa
210 Gpa
210 GPa
Fracture toughness 10 Mpa √m
7 Mpa √m
5.9 MPa √m
Compression resistance Not provided 2000 MPa Not provided
Fatigue strength (ring
on disk)
Not provided <0.002mm3/h Not provided
Corrosion resistance in
Ringer solution 37°C
Not provided <0.01(mg/m2x24h) Not provided
Flexural strength Not provided 1200 MPa Not provided
Chemical solubility
(ISO 6872)
Not provided Not provided <20 µg/cm2
TEC (25-500°C) 10.4x10-6/K 10 x10-6/K 10.5 x10-6/K
Sintering stage Presintered Densely sintered Presintered
Table 2: Overview of the framework materials
(Data provided by manufacturers)
42
MATERIALS AND METHODS
The composition of the e max Ceram® veneering material (Ivoclar Vivadent AG, Schaan,
Liechtenstein) used for the veneering of all the frameworks are listed in Table 3:
• Veneering Material • Composition
(weight%) • e max Ceram® (Ivoclar,
Vivadent)
SiO2 60-65
Al2O3 9-11
• K20 • 7-8
Na20 7-8
F 1-1.5
Zr02 1-1.5
ZnO 2-3
TiO2 1-1.5
TEC (10-6 K-1) 9.5 ± 0.25
Table 3: Composition of the e max Ceram® (emC) veneering material
(Data provided by manufacturer)
5.1.3 Materials used for the cementation procedure
KetacTM Cem MaxicapTM (3M ESPE) is a glass-ionomer cement, which consists of a powder
and a liquid. The powder is composed of glass powder and pigments. The liquid is composed
of polycarbonacids, tartaric acid, water and conservation agent.
The cement is provided in a capsule with standardized doses of powder and liquid ensuring a
controllable quality of the cement properties. The system also provides a Maxicap activator to
activate the capsules for 2 seconds; a high-frequency mixer with 4300 oscillations/min, such
as Capmix (8-15 sec mixing time), or a rotation mixer Rotomix (7-12 sec) for mixing; and a
Maxicap applier for application. The times applied for an ambient temperature of 23°C are as
follows:
Activation time: 2 sec
Mixing in Rotomix: 8 sec
Mixing in high-speed mixer: 10 sec
Working time (from beginning of mixing): 3 min
Setting time (from beginning of mixing): 7 min
43
MATERIALS AND METHODS
An overextended working time causes the loss of adhesion to enamel and dentin.
Impression and die materials
Twinduo (Picodent, Wipperfürth, Germany)
The material is an additional cross-linking high viscosity silicone impression material
presenting a 30-second mixing time, 1.5-minute working time and 6- to 7-minute setting time.
It presents 0.9% deformation under load, 99.2% elasticity after deformation and 0.1% linear
dimensional change.
Dimension Garant L (Espe, Seefeld, Germany)
This is a hydrophilic low consistency additional polymerization silicone material. It presents
4.5% deformation under load (ISO), 99.9% elasticity after deformation (ISO) and –0.20%
linear dimensional change (ISO, after 24h). The setting time is 5.5 minutes after mixing.
Firmer Set Putty (Espe, Seefeld, Germany)
The material is an additional cross-linking high viscosity silicone impression material with a –
0.05% dimensional change and a 0.3% compression set. The mixing time is 45 seconds and
the setting time is 5 minutes.
GC Fujirock (GC Belgium)
This is a Type 4 dental stone with improved physical properties and workability. The
recommended water/powder ratio is 20 ml/100gr. The setting expansion is 0.08% and the
compressive strength is 53 Mpa.
44
MATERIALS AND METHODS
5.1.4 Additional Materials (Table 4):
Material/Equipment Company
Anti-Rutsch Lack Wenco-Wenselaar GmbH, Hilden, Germany
Artificial oral environment/
Thermocycling System
Willytech, Munich, Germany
Gebrüder Haake GmbH, Karlsruhe, Germany
Steatite ceramic ball Hoechst Ceram Tec, Wunsiedel, Germany
Diamond Burs No 386.023, No
8368.023
No 837KR.012, No 8837KR.012
Gebr. Brasseler, Lemgo, Germany
Technovit 4000 Zwick, Ulm, Germany
Occlu Plus-Spray Hager & Werken GmbH &Co. KG, Duisburg,
Germany
FG-Diabolo burs Bredent, Senden, Germany
Zwick Z010/TN2S Zwick, Ulm, Germany
Dentona esthetic-base gold dental stone Detmold, Germany
Table 4: Overview of the additional materials used for the experiment
5.2 METHODS
5.2.1 Representative model
The representative model consists of two abutments (a second premolar and a second molar)
and one pontic, which functions to replace the first missing molar. In a study by Stambaugh
and Wittrock (1977), who reported that the average mesiodistal width of the mandibular first
molar is 10.92 mm, the mesiodistal width of the pontic used in the study was 11 mm.
The selected model was embedded in a sample holder of an artificial oral environment
(Wilytech, Munich, Germany) using a silicone putty material (Twinduo, Picodent,
Wipperfürth, Germany). The occlusal table of the model was set to be parallel to the
horizontal plane. A silicon mold (Twinduo, Picodent, Wipperfürth, Germany) of the model
was then fabricated covering at least 2 mm above the edge of the sample holder. This mold
was also used as the negative form for fixing the abutments in the sample holder. The
45
MATERIALS AND METHODS
abutment teeth were set in place in the mold with wax, which covered their cervical area to an
extent of 2 mm apically from the cemento-enamel junction (CEJ). The 2 mm represent the
biological width, considering that the margin of the preparation was set approximately to the
height of the CEJ.
5.2.2 Artificial periodontal membrane
It has been demonstrated that abutment mobility has a big influence on the fracture resistance
of FPDs (Kelly et al, 1995). Therefore, in order to imitate the physiological tooth mobility, all
roots were covered with an artificial periodontal membrane made out of a gum resin (Anti-
Rutsch-Lack, Wenko-Wenselaar GmbH, Hilden, Germany) (Kern et al, 1993; Koutayas et al,
2000). This resin covered the root starting 2 mm apically from the CEJ to the root tip
(biological width principle).
5.2.3 Embedding models in the sample holders
After the gum resin has dried, the silicon mold with the abutments in place was attached to the
sample holder, which was previously isolated with Vaseline (Weises Vaselin Lichtenstein,
Winthrop, D-Fürstenfeldbruck). A self-curing polyester resin (Technovit 4000, Kulzer,
Wehrheim, Germany) was then mixed and poured into the sample holder. After the resin had
set, the silicon mold was removed and the abutments were cleaned. The resin specimens were
then stored in 0.1% thymol solution.
5.2.4 Tooth preparation (Fig. 2)
The abutment teeth were prepared with a 1.2-mm circular deep chamfer, an occlusal reduction
of 1.5 mm, a facial reduction of 1.2 mm and a convergence angle of 6°. The heights of the
premolars and the molars were 6 mm and 5 mm, respectively. All transitions from the axial to
the occlusal surfaces were rounded. Homogeneous and smooth surfaces were achieved.
46
MATERIALS AND METHODS
Figure 2: Representative tooth preparation
5.2.5 Impression procedure
The impression technique used is known as the simultaneous Putty/Wash-Technique. First, a
thin coat of 3M ESPE tray adhesive was brushed on the mini-tray and left to air dry for a
minimum of 5 minutes. After the prepared abutments were thoroughly dried, Dimension
Garant L was placed around them using a syringe. Equal volumes of Firmer Set Putty base
and catalyst were then mixed, put on the mini-tray and placed parallel to the tooth line. It was
held in position without pressure until the material was set (5 minutes). The impression was
then removed from the abutments, and after one hour the master model was poured.
5.2.6 Fabrication of master models
In order to avoid introduction/encapsulation of bubbles into the master model, the impressions
were first conditioned with a silicone surfactant. Then, GC Fujirock type 4 dental stone was
added to the water within 10 seconds and mixed uniformly for 60 seconds by mechanical
spatulation under vacuum. The recommended water/powder ratio is 20ml/100gr. After a
setting time of at least 45 minutes, the master model was removed from the impression.
Exceptionally, for the fabrication of the Vita In Ceram® 2000 YZ Cubes® master models, a
special scannable dental stone (Dentona esthetic-base gold, Detmold, Germany) was used.
47
MATERIALS AND METHODS
5.2.7 Fabrication of all-ceramic FPDs
5.2.7.1 Manufacturing the framework
The frameworks of all groups had a connector height of 3 mm, a connector width of 3 mm, a
0.7 mm occlusal thickness and an axial thickness of 0.5 mm.
● Procera Zirconia®, Nobel-Biocare AB, Göteborg, Sweden (Fig. 3)
Figure 3: Procera® Zirconia frameworks
After the fabrication of the stone die, a tactile scanner (Procera Forte, Procera, Nobel Biocare,
Göteborg, Sweden) was used to scan the dies in order to duplicate the shape and to identify
the margins with precision. This scanner is indicated not only for single units but also for
bridges with their soft tissues, neighboring teeth, and bite registrations. The whole scanning
process takes from 6-10 minutes. The data were transmitted to the Procera Software CAD-
Application, where the design of the restoration took place. A 0.7 mm occlusal thickness, 0.5
mm circumferential thickness and 9 mm2 connector surface were taken into consideration for
the design of the frameworks. Then, the computer-designed framework was oversized by 20%
to compensate for contraction which occurs during the final sintering. After completion of the
design-procedure, the enlarged dies were fabricated. Zirconia is dry-pressed against the
enlarged dies, and the temperature is raised to a temperature similar to the presintering stage.
At this point, the enlarged and porous copings are stable. Their outer surfaces were milled to
the desired shape and the copings, now removed from the enlarged dies, were placed into the
furnace and fired to full sintering. During this cycle, the copings shrank to fit the dimensions
of the original working dies. Subsequent fitting adjustments were made after sintering using
48
MATERIALS AND METHODS
rotary diamond instruments with water cooling (FG-Diabolo burs, Bredent, Senden,
Germany). The completed copings were then ready for the veneering process.
● DC-Zirkon (DCS President System, Allschwill, Switzerland) (Fig. 4)
Figure 4: DC-Zirkon® frameworks
The DCS President® system consists of three main parts: I) the Preciscan®, a fully automatic
laser scanner, II) the DCS Dentform® software and III) the Precimill® machining center. The
Preciscan® laser scanner measures preparation dies (stone dies) as well as the whole cast (up
to 14 single abutments or entire FPDs). After measuring, the collected data were then
digitized and transmitted to a computer where the FPD framework was designed and
calculated (CAM process). Data on the framework were subsequently forwarded to the
Precimill® machining center, where a framework (0.5 mm axial thickness, 0.7 mm occlusal
thickness and 3x3 mm connector dimensions) was milled out of a densely sintered zirconia
block. This ZrO2 block was manufactured under optimized industrial conditions by the
sintering and hot isostatic post-compaction (HIP) of tetragonal zirconia polycrystals (TZP)
(TKT Metoxit® AG, Thayngen, Switzerland). With this system there was no need to confront
any dimensional changes because the material was already sintered. Subsequent fitting
adjustments were made using rotary diamond instruments with water cooling (FG-Diabolo
burs, Bredent, Senden, Germany), and the frameworks were ready for further veneering.
49
MATERIALS AND METHODS
● Vita In Ceram® 2000 YZ Cubes® (Vita Zahnfabrik, Bad Säckingen, Germany) (Fig. 5)
Figure 5: Vita In Ceram® YZ Cubes® frameworks
Dies of the Vita group were scanned using the Cerec InLab scanner (Sirona, Bensheim,
Germany). The data were then digitized and transmitted to a computer where the framework
was designed and calculated. Presintered Y-TZP blocks (Vita In Ceram® 2000 YZ Cubes®,
Vita Zahnfabrik, D-Bad Säckingen) were milled in the CAM unit (Cerec InLab, Sirona,
Bensheim, Germany). The unit mills an enlarged framework (0.5 mm axial thickness, 0.7 mm
occlusal thickness and 3x3 mm connector dimensions) out of a presintered block to
compensate for the later sintering shrinkage. The milled frameworks were then carefully
removed from the machine and separated from the block holders at the milled side using
diamond cutting instruments. The frameworks were subsequently postsintered using a high
temperature furnace (Vita ZXrcomat, Vita Zahnfabrik, Bad Säckingen, Germany). The
temperature in the firing chamber would not exceed 1600° C. Subsequent fitting adjustments
after sintering were made using rotary diamond instruments with water cooling (FG-Diabolo
burs, Bredent, Senden, Germany).
5.2.7.2 Veneering procedures (Fig. 6)
All frameworks of the test groups were veneered with the same ceramic veneering material (e
max Ceram, Ivoclar-Vivadent, Schaan, Liechtenstein) in the Programat 100 oven (Ivoclar-
Vivadent, Schaan, Liechtenstein). Silicon keys were made in order to control the thickness of
the veneering material. The veneering process of all three types of zirconia-frameworks with
e max Ceram® is listed in Table 5:
50
MATERIALS AND METHODS
Procedure Preheating
(°C)
Heating
rise (min)
Vacuum
in (°C)
Vacuum
out (°C)
Holding
time
(min)
High
temp.
(°C)
Heating
rate
(°C/min
)
Opaque 403 6 450 799 2 800 50
Dentin 403 6 450 759 2 760 50
Glaze 403 6 450 729 1 725 50
Table 5: Firing chart of the e max Ceram® (emC) (Ivoclar-Vivadent, Schaan,
Liechtenstein)
Figure 6: Procera Zirconia® framework veneered with the e max Ceram® veneering
5.2.8 Cementation of the FPDs
PE, Seefeld, Germany) glass-ionomer cement was used
t
material
The KetacTM Cem MaxicapTM (3M ES
for the cementation of all FPDs. The abutments were thoroughly cleaned with 0.1%
chlorhexedine and air-dried so that the surface has a matte shiny appearance. Similarly, the
inner surfaces of the copings of the FPDs were cleaned with chlorhexedine 0.1% and
completely air-dried. A capsule of KetacTM Cem MaxicapTM was then activated for 2 sec in
the Maxicap activator and further mixed for 10 sec in the Ro omix. The capsule was placed in
51
MATERIALS AND METHODS
the Maxicap applier and further applied on a glass-slab. Afterwards, the cement was evenly
applied on the inner surfaces of the copings with a small brush and the FPDs were placed on
the abutments and held in place under finger pressure for 2-3 minutes. After it was set (7
minutes), excessive cement was removed using a scalpel.
5.2.9 Dynamic loading of the test samples
The artificial mouth (Figure 7) allows the evaluation of dental restorative systems under
clinically relevant conditions (Krejci et al, 1990). The artificial oral environment (Willytec,
Munich, Germany ) consists of 8 identical sample chambers, two stepper motors controlling
the vertical and horizontal movement of the samples against an antagonist, a hot and cold
water circulation system (Haake, Karlsruhe, Germany) and a computerized control unit (Kern
et al, 1999). Half of the samples of each group were subjected to 1.2 million chewing cycles
by a reproducible dynamic occlusal load, which corresponds to five years of clinical function
(Krejci et al, 1990). The applied load was 50 N (De Boever et al, 1978; Krejci et al, 1990),
and the thermocycling was 5°C to 55°C for 60 seconds each, with an intermediate pause of 12
seconds, maintained by the thermostatically-controlled liquid circulator (Haake, Karlsruhe,
Germany) (Fig. 8). A 6-mm diameter ceramic antagonist Steatit® ball (Höchst Ceram Tec,
Wunsiedel, Germany) was applied vertically onto the occlusal surface of the pontic of the
FPDs. During dynamic loading, all samples were examined twice a day and any fracture of
the teeth or the porcelain was recorded as a failure.
Cold/hot bath temperature 5° C/55° C
Dwell time 60 sec
Vertical movement 6 mm
Horizontal movement 0.5 mm
Descending speed 60 mm/s
Rising speed 55 mm/s
Forward speed 60 mm/s
Backward speed 55 mm/s
Applied weight per sample ) 5 kg (49 N
Cycle frequency 1.6 Hz
52
MATERIALS AND METHODS
Table 6: Test parameters of the artificial mouth
ht sample chambers
oval of cold water,
) pump for removal of warm water, (8) pump for application of cold water, (9) pump for
pplication of warm water, (10) motor table, (11) table
Figure 7: Schematic drawing of the dual-axis chewing simulator with eig
(Willytech, Munich, Germany) (Kern, 1999)
(1) upper crossbeam, (2) lower crossbeam, (3a) water reservoir (in), (3b) water reservoir
(out), (4) filter for cold water, (5) filter for warm water, (6) pump for rem
(7
a
53
MATERIALS AND METHODS
Figure 8: Schematic drawing for one chewing chamber (Kern, 1999)
54
MATERIALS AND METHODS
5.2.10 Survival rate
uring the dynamic loading, all samples were examined twice a day. Fractures of the tooth or
e porcelain were recorded as failure.
.2.11 Fracture resistance test
ples of all test groups were loaded until fracture occurred using a universal-testing
achine (Zwick Z010/TN2S, Ulm, Germany). Tin foil of 1mm thickness (Dentaurum,
as placed over the occlusal surface of the first molar (pontic) to
r a stroke control of 2 mm/min. The loads required to
acture the samples were recorded with the Zwick testXpert® V 7.1 software.
.2.12 Statistics
D
th
5
All sam
m
Ispringen, Germany) w
achieve a homogeneous stress distribution. A perpendicular load was applied to the occlusal
surface of the first molar (pontic), unde
fr
5
The statistical analysis of the fracture resistance tests was performed by Dr T. Gerds, Institute
of Medical Biometry and Medical Informatics, Albert Ludwigs University, Freiburg,
Germany. Fracture resistance data were analyzed by nonparametric ANOVA using Kruskal-
Wallis and Wilcoxon rank sum tests (R Development Core Team 2005, R: A language and
environment for statistical computing, R Foundation for statistical computing, Vienna,
Austria. ISBN 3-900051-07-0. http://www.R-project.org) with a significance level of 0.05.
oxplots were used for visualizing the data.
B
55
RESULTS
6 Results
.1 Survival rate of all-ceramic FPDs after aging
ll specimens subjected to aging survived 1,200,000 cycles of dynamic loading. No chipping
f the veneering ceramic or decementation of the FPDs was recorded.
.2 Fracture resistance tests (Table 7, Fig. 9)
he smallest fracture resistance value was observed in the aged Procera group (1044 N),
value was observed in the non-aged DCS group (1993) (Table 7). The
highest median fracture resistance value without aging occurred in the Vita CerecInLab group
(1702 roups. After aging, the
ighest value was found in the DCS group (1618 N) followed by the Vita CerecInlab (1556
Group Minimum 1st Median Mean 3rd
Quartile
Maximum
6
A
o
6
T
whereas the highest
N), followed by the DCS (1683 N) and Procera (1522 N) g
h
N) and Procera (1256 N) groups.
Quartile (± S.D.)
Procera
Zirconia
initial
1105 1277.588 1522 1496±260 1727.325 1800
Procera
Zirconia
aged
1044 1130.75 1256 1297±242 1347.707 1783
DC-Zirkon
Initial 1278 1489.378 1683 1659±245 1845.37 1993
DC-Zirkon 1175 1522.432 1618 1580±197 1713.403 1804
aged
Vita
CerecInLab 1472 1658.888 1702 2 1776.035 1946
initial
1713±14
Vita 1394 1478.702 1556 1593±174 1663.438 1854
56
RESULTS
CerecInLab
aged
Table 7: Results o racture e of di ups u stance
s of individual samples are listed in the appendix):
f the f resistanc values fferent gro in N (fract re resi
value
smaller fracture resistance of Procera compared to Vita CerecInLab (Wilcoxon test: p=0.015)
and to DCS (Wilcoxon test: p=0.038).
Fig. 9 Box plot of the results of the fracture resistance test in N
The central box shows the data between the 1st-Quartile and the 3rd-Quartile, the median is
represented by a line.
Compared to the values at the initial stage, artificial aging reduced the fracture resistance by
4.76%, 13.3% and 7% for groups DCS, Procera and Vita CerecInLab, respectively. This
reduction, however, was not statistically significant for any of the groups tested. Similarly, no
significant differences were found for the fracture resistance comparisons between different
groups before artificial aging (Kruskal-Wallis test: p=0.3). After artificial aging, the Kruskal-
Wallis test showed a significant group effect (p=0.03). Wilcoxon test revealed significantly
57
RESULTS
6.3 Fracture patterns
The location and mode of failure of the specimens after the load-to-fracture test are
summarized in Table 8. Most of the fractures before and after aging occurred either at the
distal connector or at both connectors.
able 8: Location and mode of failure after the load-to-fracture test
Group Abutment
fracture
Distal
connector
Mesial
connector
Both
connectors
DCS/initial 1 7 - -
DCS/aged - 4 - 4
Procera/initial - 2 1 5
T
Procera/aged - 8 - -
Vita
CerecInLab/initial - 4 3 1
Vita
CerecInLab/aged - 3 - 5
58
RESULTS
the group without artificial aging, 5 FPDs fractured at both connector sides; 1 FPD at the
mesial connector (premolar side) and 2 FPD at the distal connector (molar side) followed by
decem as well.
In the group after artificial aging, all 8 FPDs fractured at th
Dec as observed 2 specimens.
Although it was difficult to assess whether the fractures started at the loading point or at the
conn ractures were perpendicular to mesial-distal eworks in a
smooth curve between the loading point and the gingival side of the connector (Fig. 10)
6.3.1 Procera Zirconia group
In
entation
e distal connector (molar side).
ementation w in
ectors, the f the axis of the fram
igure 10 Representative figure of the fracture pattern of Procera group. F
59
RESULTS
e), 3
PDs showed fractures at both connector sides and 1 FPD showed a fracture of the self-curing
sin (Technovit 4000, Zwick, Ulm, Germany) at the premolar region combined with
onnector fracture at this area.
Similar to the Procera group, the fractures were perpendicular to the mesial-distal axis of the
frameworks in a smooth curve between the loading point and the gingival side of the
connector (Fig. 11).
6.3.2 DCS group
In the group without artificial aging, 7 FPDs fractured at the distal connector (molar side) and
1 FPD showed a fracture of the premolar tooth combined with a total fracture at the distal
connector (molar side)
In the group after artificial aging, 4 FPDs fractured at the distal connector (molar sid
F
re
c
ig. 11 Representative figure of the fracture pattern of DCS group F
60
RESULTS
ng, 3 FPDs fractured at both connector sides, 3 FPDs at the
s in the previous test groups, it was difficult to determine the origin of the fracture. The
actures were perpendicular to the mesial-distal axis of the frameworks in a smooth curve
etween the loading point and the gingival side of the connector (Fig. 12).
6.3.3 Vita CerecInLab group
In the group without artificial aging, 3 FPDs demonstrated fractures at the mesial connector
(premolar side), 4 FPDs at the distal connector side (molar region) and 1 FPD fractured at
both connector sides.
In the group after artificial agi
distal connector side (molar region) and the rest 2 showed fractures of the self-curing resin
(Technovit 4000, Zwick, Ulm, Germany) around the premolar abutments, in combination
with fractures at both connector sides.
A
fr
b
ig. 12 Representative figure of the fracture pattern of the Vita CerecInLab group. F
61
DISCUSSION
7 Discussion he design and tests conducted in this study were chosen to better simulate clinical
metals were used for testing, such as Ni-Cr-Fe alloy (Wiskott et al, 1996), stainless steel
(Wilson, 1994), and brass (Yamashita et al, 1997b). Their advantages are that the metal
abutments have identical physical properties and dimensions. However, the elastic and
bonding properties of these abutments cannot be compared to those of natural teeth.
Plastic abutments made out of composite or epoxy resin have been also used as abutments for
fracture resistance testing of all-ceramic crowns (Scherrer and de Rijk, 1993; Yoshinari and
Derand, 1994; Neiva et al, 1998) and post and core systems (Schmeißner, 1977; Wegmann,
1987). Their advantages are similar to those of metal abutments, including a modulus of
horic acid
eiva et al, 1998), but they lack bonding characteristics as they do not consist of water and
rganic substances. In a comparative in vitro study, Rosentritt et al (2006) investigated the
fluence of the abutment material, using human, polymer and alloy abutments, on the
ramic FPDs. They found out that the application of the alloy
T
conditions. The number of the specimens tested and the use of water rather than artificial
saliva during testing are limitations that may affect the interpretation of the results.
7.1 Methods
7.1.1 The use of natural teeth as abutments
In this study, extracted human teeth were used as abutments because their modulus of
elasticity, bonding characteristics, thermal conductivity and strength are closer to the clinical
situation than those of metal, plastic and animal teeth. Human teeth have also been used in
other studies (Dietschi et al, 1990; Haller, 1990; King and Aboush, 1999; Chitmongkolsuk et
al, 2002; Rosentritt et al, 2006). The extracted human teeth were stored in 0.1% thymol
solution, preventing them from drying out and thereby becoming brittle (Helfer et al, 1972),
and also inhibiting microbial activity (Sparrius and Grossman, 1989).
In several studies, metal abutments were used to test the fracture strength of bridges (Kappert,
1990; Bieniek, 1994; Ludwig, 1994; Kappert, 1995; Rosentritt et al, 2006). Different types of
elasticity similar to that of human dentin. They can be etched with 34% phosp
(N
o
in
fracture resistance of all-ce
62
DISCUSSION
abutments clearly lead to an overestimation of the fracture resistance of the ceramic FPDs.
Human and polymeric abutments showed similar influence, and the combination with an
artificial periodontium lead to a 70% reduction of the fracture resistance after artificial aging.
Bovine teeth have similar bonding characteristics, modulus of elasticity and tensile strength to
human al, 1994). They have also been used for testing the fracture strength of
eramic restorations (Mesaros et al, 1994), the bonding strength of luting materials
and the fracture resistance of different
screpancy between bovine and
uman teeth, however, the use of bovine teeth for testing the fracture resistance of all-ceramic
hewing simulation and
ble aging
teeth (Sano et
c
(Phrukkanon et al, 1998; Hosoya and Tominaga, 1999)
post systems (Isidor et al, 1999). Because of the great size di
h
FPDs is difficult.
7.1.2 Artificial periodontal membrane
In the present study, a thin layer of gum resin (Anti-Rutsch-lack®) was painted on the roots of
the abutments to imitate physiological tooth mobility during both c
fracture resistance testing. Kern et al (1993) and Koutayas et al (2000) showed that it can
mimic tooth mobility similar to physiological tooth movement (Mühlemann, 1951).
Abutment mobility has been demonstrated as a decisive factor in the evaluation of fracture
resistance, and when a small abutment rotation is allowed, failure is more likely to occur
(Kelly et al, 1995). A series of in vitro studies tested the fracture resistance of In Ceram
Alumina FPDs using variable materials to simulate the abutment mobility and varia
methods. The results after chewing simulation, with and without artificial periodontal
membrane, were 523 N and 337 N (DeLong and Douglas, 1983), 676 N and 256 N (Rosentritt
et al, 2000) and 919 N and 305 N (Scherrer et al, 1996), respectively. Similarly, after
thermocycling in artificial saliva, the values obtained were 2225 N for groups without
artificial periodontium versus 703 N for groups with periodontal membrane (Kappert et al,
1991). These observations are in agreement with the reports of another recent study showing
that the fracture strength of Empress II® FPDs (Ivoclar-Vivadent, Schaan, Liechtenstein) was
significantly influenced by the use of an artificial periodontium device (1 mm Impregum, 3M
Espe, G-Seefeld)(Rosentritt et al, 2006). Resigning on the periodontium during artificial
aging caused fracture strength values almost twice as high as the ones with periodontium.
More distinct was the influence of the artificial periodontium by the non-aged FPDs: rigid
teeth showed results that were three-times higher than those with the polyether layer.
63
DISCUSSION
7.1.3 The antagonistic material
In this in vitro study, ceramic balls (Steatite, Hoechst Ceram Tec, Wunsiedel, Germany) were
used as antagonists during the artificial aging in the chewing machine. These antagonists have
Vickers hardness similar to that of enamel. Few studies have investigated the influence of the
ntagonistic material on the fracture resistance of FPDs. In an in vitro study, the use of human
antagonists instead of ceramic balls significantly decreased the loading capacity of FPDs
y, the antagonistic material and design showed
similar effect on the FPDs, reducing insignificantly the fracture strength by 175 N
ult von
teyern et al, 2001). Other studies showed that the stress concentration on the connectors of
e FPDs is reduced with a connector of at least 4 mm in height (Kamposiora et al, 1996;
Pospiech et al, 1996). A study using strain gauges in posterior FPDs indicated that the strain
distribution in-vivo is different from that in-vitro. In-vitro, the marginal portion under the
a
(Condon and Ferracane, 1996). In another stud
a
(Rosentritt et al, 2006).
7.1.4 Preparation design and connector dimensions
Although there are no standards in the literature in terms of preparation design for zirconia-
based all-ceramic FPDs, it is recommended that an occlusal reduction of 1.5-2 mm , a facial
reduction of 1-1.5 mm, a deep chamfer or circular shoulder with inner rounded angle between
1.0 and 1.5 mm wide, and an occlusal angle of convergence not greater than 10 degrees be
employed (Friedlander et al, 1990; Chiche, 1994; Shillingburg, 1997; Sorensen et al, 1998;
McLaren and White, 1999; McLaren, 2000).
In the present study, a circular 1.2-mm deep chamfer preparation with a 6°-convergence
angle, 1.5-mm occlusal reduction and smooth, rounded transitions from the axial to the
occlusal surfaces was the aim. The increased fracture resistance of the new high-strength
ceramics compensates for the decreased thickness of the material required.
The maximum strain in FPDs is located in the connector area. In the present study, all
connectors had 3x3-mm well rounded dimensions. Variable connector dimensions have been
suggested in the literature: 4x4 mm (Kappert, 1990), 3x3 mm (Futterknecht, 1990; McLean,
1993), 3 mm buccolingually and 4 mm occlusogingivally (McLaren, 1998) and a 4 mm
diameter in the case of a molar replacement and 3 mm diameter for other cases (V
S
th
64
DISCUSSION
cusp during loading was greatly strained. In-vivo, however, the whole portion of the FPD
trained, and the strain values were greatest on the buccal and lingual portions of the posterior
is limited and the biologic width must be respected (Sorensen et al,
999). Because the core ceramic is significantly stronger than the veneering porcelain, it may
be advisable in the case of a FPD not to veneer the core material at the tissue side of the
crucial (Sorensen et al, 1999;
cLaren and White, 2000; Guazzato et al, 2004d). In contrast to the former study, Sundh et
he
e
materials must be
not valid. The ring-on-ring biaxial bend test offers a larger specimen area or volume subjected
s
retainer, and on the distal connector (Yamashita et al, 1997a). Enlarging the connector cross-
section reduces the peak stresses and hence the probability of fracture for conventional
systems. Oversized connectors, however, limit the clinical range of use, compromise plaque
control and lead to unaesthetic results. Under clinical conditions, the occlusal contact and the
gingival tissue define the limits of the connector dimensions. If the minimum vertical
dimension required for the connectors is not available, the clinician may consider performing
electrosurgery to remove the soft tissue to gain space for the connector height, although the
extent of tissue removal
1
connectors and in areas where esthetic considerations are not
M
al (2005) showed no significant difference in fracture resistance of Y-TZP FPDs between t
veneered specimens and those heat-treated without veneering. As the few available studies ar
controversial, further studies are necessary to investigate the design and influence of the
environment of restorations with the core material exposed (Guazzato et al, 2004c).
7.1.5 Clinical relevance of fracture resistance tests
In order to assess the suitability of new experimental ceramic materials for dental application,
a number of in-vitro studies testing the properties and behavior of the
carried out. Taking into consideration the brittleness of ceramics, the fracture strength of
experimental ceramic materials should be comparable to those of other accepted materials to
validate their use for further clinical testing. Fracture strength tests of ceramic materials are
important for determining or maximizing the expected lifetime with an acceptable low
probability of failure (Ritter, 1995b).
Considering the practical difficulties in the preparation of specimens, the most commonly
used testing methods for ceramic materials are four-point bending, three-point bending, and
biaxial bending; which include the ring-on-ring (Kao, 1971), ball-on-ring (McKinney, 1970),
and piston-on-three-ball tests (Kirstein, 1967). It is important to note that the failure stresses
derived from different testing methods are significantly different, and a direct comparison is
65
DISCUSSION
to the maximum stress compared to the three-point bending and the piston-on-three ball tests.
It is also unaffected by edge failure, and resembles the clinical condition as it generates the
greatest number of interfacial failures (Zeng et al, 1996; Guazzato et al, 2004c). Furthermore,
the disk-shaped specimens are preferable because they have an area similar to dental
restorations (Ban and Anusavice, 1990; Thompson, 2000).
Thompson (2000) investigated the effect of testing method and variation of the core-thickness
ratio of bilayered ceramics on the mode and origin of failure, and found a positive effect. In
this study, 250 out of 270 specimens delaminated; and for these testing configurations and
relative layer heights, the flexural strength of the core material generally proved greater than
interface toughness. In regard to the origin of failure, none of the “clinically similar”
specimens (core/veneer ratio 1:2) failed at the interface with any testing method. When the
ratio becomes 1:2, which resembles the geometry of a FPD connector, the interface failure
increases to 43%. This is in agreement with the findings of Kelly et al (1995), who reported
that the most interface failure origins (failure rate 70-78%) occurred at the gingival side of the
FPD connectors. Another study, carried out by Wakabayashi and Anusavice (2000),
investigated the effect of the core/veneer thickness ratio and of the elastic modulus of the
substrate on the crack initiation of bilayered ceramic disks. They found that the core/veneer
ickness ratio is the dominant factor that controls the failure initiation site (shifts from the
veneer to the core as the ratio increases) in bilayered ceramics; but the increase in elastic
dditional factors that need to be considered when trying to compare the results gained from
ceramic, the test environment must be the same as the
th
modulus did not affect the crack initiation site.
A
different studies are the use of mobile or immobile abutments, the static or dynamic mode of
load, and the wet or dry environment. It has been shown that the values obtained when using
immobile abutments are higher than those when using mobile abutments (Tinschert et al,
1999), whereas dynamic loading reduces the flexural strength of dental ceramics compared to
their static values (Geis-Gerstorfer and Fässler, 1999). Furthermore, water and temperature
changes reduce the fracture strength of ceramics. For the strength to accurately reflect the
variability and time dependency of a
service environment, and the strength-controlling flaw population must be the same as that
responsible for failure in service. Therefore, it is generally recommended that test samples and
mode of loading be chosen to closely simulate the actual components in service (Ritter,
1995a, b; Kelly, 1999). Thus, a direct comparison between the results obtained in the different
studies is difficult, as is extrapolation of in-vitro results to clinical situations.
66
DISCUSSION
7.1.6 Clinical relevance of the artificial aging process
In the oral environment, the forces applied on dental restorations are of a cyclic nature
(Lundgren and Laurell, 1984). In addition, the presence of moisture and temperature changes
e chewing machine fulfils the parameters
leads to slow flaw propagation and degradation of the mechanical properties of dental
ceramics (Kelly, 1999). Therefore, instead of monotonic static loading, it is more clinically
relevant to test the specimens under fatigue load in a chewing simulator (DeLong and
Douglas, 1983; Krejci et al, 1990; Kelly, 1999). Various chewing simulators have been used
in several in vitro studies to simulate clinical conditions and evaluate dental restorative
systems under clinically relevant conditions (DeLong and Douglas, 1983; Krejci et al, 1990;
Strub and Beschnidt, 1998; Behr et al, 1999; Kern et al, 1999; Koutayas et al, 2000). After
100,000 fatigue loading cycles, the flexural strength for In Ceram, IPS Empress and Dicor
was approximately 50% of the initial strength (Schwickerath, 1996).
The chewing simulator used in this study was developed to reproduce the in vivo
environment by adding moisture and a controlled temperature to the test conditions (Kern et
al, 1999). The artificial chewing cycle in the artificial oral environment is designed to
correspond to physiologic conditions. The magnitude, duration and frequency of the force
applied are comparable to values reported in the literature (Bates et al, 1975, 1976; Bradley,
1996). Krejci et al (1990) indicated that th
concerning chewing motion and thermal changes reported in the literature. Fatigue produced
by 240,000-250,000 cycles in the chewing simulator corresponds to a period of 1 year of
clinical service. Therefore, 1,200,000 chewing cycles correspond to a five-year clinical
service (Kern et al, 1999).
In a recent study, the influence of different simulation parameters on the properties of dental
restorations was compared (Rosentritt et al, 2006). It was found that the duplication of
chewing frequency, lateral sliding or continued force increased during aging did not
significantly influence the fracture resistance of FPDs. A temperature gradient of 50°C led to
a significant reduction of fracture resistance (400 N), compared to constant cycling of water
with 25°C. Duplication of the mouth opening distance from 2 to 4 mm did not significantly
influence the fracture resistance. Increasing the loading force from 50 to 150 N significantly
reduced the loading capacity. A staircase load-increase to 150 N showed no significant
differences compared to a test with a constant load of 150 N.
67
DISCUSSION
7.2 Results
7.2.1 Survival rate after the chewing simulation
In the present in vitro study, one-half of all samples were exposed to the artificial oral
environment to simulate a five-year clinical service, before the fracture resistance test was
performed. All the exposed samples survived, exhibiting no fractures or chipping off defects.
7.2.2 Fracture resistance tests
To date, conventional porcelain-fused-to-metal fixed partial denture (PFM) is still considered
the gold standard, in terms of predictability, for the rehabilitation of edentulous spaces at the
alues (before and after artificial aging) were much lower than the
abutments, and reported mean
alues of 1300 N after exposure to dynamic loading (Rosentritt et al, 2003). The fracture
resistance testing of Vita CerecInLab® (Vita CerecInLab Zahnfabrik, Bad Säckingen,
Germany) posterior FPDs, also investigated in the present study, revealed values ranging from
posterior region. Therefore, the resistance of PFMs may serve as a guideline for the new
tested ceramic materials. A direct comparison among the studies, however, seems reasonable
only if the test methodology is the same.
Chitmongkolsuk et al (2002), following the same test parameters with the ones in the present
study, reported fracture resistance values for posterior PFMs of 3500 and 2800 N, before and
after artificial aging, respectively. In the present study, the mean fracture resistance values for
Procera were 1496 N and 1297 N, for DCS 1659 N and 1580 N and for Vita CerecInLab 1713
N and 1593 N, before and after artificial aging, respectively. The artificial aging lead to a
degradation of the fracture resistance behavior in all groups tested, but this influence was not
stastically significant. All v
corresponding ones reported for PFMs (3500, 2800 N) but still much higher than 1000 N,
which is considered as the minimum demanded resistance of a material for application at the
posterior region (Schwickerath, 1986, 1994; Geis-Gerstorfer and Fässler, 1999; Marx et al,
2001; Tinschert et al, 2001a). The fracture resistance of Y-TZP-based all-ceramic FPDs was
evaluated in several in vitro studies. Tinschert et al (2001a) reported mean values of 2289 N
(veneered) and 1900 N (only substructure) for DC Zirkon-based (DCS President System,
Allschwill, Switzerland) FPDs, when using fixed metal abutments and without dynamic
loading processes. Another research group investigated the fracture resistance of “green”
stage Y-TZP Lava® (3M ESPE, G) FPDs, placed on human
v
68
DISCUSSION
900 N (only substructure) to 1900 N (veneered) when using fixed metal dies, 3x3 mm
onnector dimensions and exposure to dynamic loading (Sundh and Sjogren, 2006).
Following a like experimental design, the same study group reported values for HIPed Y-TZP
ing from 3291 N (only substructure)
2237 N and 1973 N, depending on the veneering material applied (Sundh et al, 2005).
.2.3 Influence of the chewing simulation on the fracture resistance
nt (P>0.05) for all groups. This outcome is in
c
Denzir (Cad.esthetics AB, Skelleftea, Sweden) FPDs rang
to
Because of the variety of the experimental designs followed, including connector dimensions,
abutment selection/mobility and exposure to fatigue loading, a direct comparison between the
values reported in previous studies may be misleading.
7
In the present study, the exposure to fatigue loading reduced the fracture resistance of all
tested groups. The Procera group showed a 13.30% reduction of the fracture resistance,
followed by 7% for the Vita CerecInLab group and 4.76% for the DCS group. This reduction,
however, was not statistically significa
agreement with previous findings showing that cyclic loading in water did not significantly
affect the fracture resistance of Y-TZP Denzir (Cad.esthetics AB, Skelleftea, Sweden) FPDs
(Sundh et al, 2005). In addition, Y-TZP ceramics have previously been identified to be high
strength materials, which were not influenced by repeated loading of up to 3000 N (Jung et al,
2000). However, when the stresses induced by repeated loading do not initially exceed the
flexural strength, the growth of subcritical flaws may result in the occurrence of delayed
catastrophic failure following 5x105 cycles (Chevalier, 1999b; Jung et al, 2000; Rauchs,
2001).
Variations in the reduction rate of fracture resistance between different groups may be
explained by different fabrication techniques for each system. Milling a framework out of a
presintered Y-TZP blank and subsequently postsintering will probably produce surface flaws
and residual/compressive stresses different than those produced in a Y-TZP framework milled
out of a fully sintered blank (Tinschert et al, 2004a; Deville et al, 2005; Sundh et al, 2005;
Chevalier, 2006; Deville et al, 2006). This will in turn lead to differences in the low
temperature degradation resistance between different systems (Chevalier, 2006).Other
comparative studies, which employ surface quality and structure detecting methodologies
such as X-Ray Diffraction (XRD), Atomic Force Microscopy (AFM) and Optical
Interferometer (OI), may elucidate the effect of different fabrication techniques on the aging
sensitivity of Y-TZP ceramics.
69
DISCUSSION
Moisture contamination has been identified to be detrimental to the fracture resistance of
ceramic-based dental restorative materials, routinely resulting in a 20% decrease in mean
fracture strength (Sherrill, 1974; Morena et al, 1986; Addison et al, 2003). Schwickerath
(1986) reported that the fracture strength of ceramic specimens was lowered by approximately
50% after 106 cycles. Similarly, combined thermal and mechanical loading significantly
ded that significant artificial aging,
ombining thermal cycling with mechanical loading, should be performed to obtain clinically
relevant results. In the literature, however, there is some controversy regarding the influence
r, 1995). In
n in vitro study, it was found that artificial aging lead to a statistically significant decrease of
dh and Sjogren, 2006). In the literature, however, conflicting
reduced the fracture resistance of Empress II® FPDs from 1832 N to 410 N (Rosentritt et al,
2006). In the same study, after testing the influence of diverse stress parameters on the
fracture resistance of all-ceramic FPDs, it was conclu
c
of moisture on Y-TZP ceramics (Dauskardt, 1987; Shimizu et al, 1993; Chevalie
a
the fracture resistance of 3-unit CAD-CAM Lava® FPDs, but not 4-unit FPDs (Rountree et al,
2001).
In this study, even after chewing simulation, all test specimens of different groups exhibited
fracture resistance values in the range of 1297 to 1593 N, which are far higher than the
minimum expected loading capacity of 1000 N for application in the posterior region. Thus,
the results indicate that Y-TZP ceramic is a material to be considered for all-ceramic FPDs in
premolar and molar regions.
7.2.4 Influence of the veneering process on the fracture resistance of zirconia-based frameworks
In order to improve their esthetic appearance, milled frameworks are veneered with
compatible veneering materials. During the veneering procedure, the frameworks are exposed
to moisture at relatively high temperatures. Y-TZP ceramic, however, are unstable over time
due to the spontaneous transformation of the tetragonal phase (t) into the monoclinic phase
(m), leading to mechanical property degradation (Chevalier, 1999a; Piconi and Maccauro,
1999). Since the t-m transformation is affected by temperature and vapor, the possibility that
the mechanical properties of Y-TZP ceramics are affected during veneering cannot be
excluded. Additionally, it has been suggested that grinding by machining introduces residual
compressive stresses on the surface, which influences the mechanical properties of zirconia
ceramics, and that subsequent heat treatment/veneering relaxes these residual stresses (Reed,
1977; Kosmac et al, 1999; Sun
70
DISCUSSION
results are reported regarding the effects of surface treatment on the strength of dental
ceramics (Zhang et al, 2004; Guazzato et al, 2005b; Sundh et al, 2005). In a study of Zhang et
al (2004), sandblasting of Y-TZP and alumina ceramics significantly reduced their strength,
while in another study sandblasting increased the strength of Y-TZP ceramics (Guazzato et al,
2005b). Moreover, the fracture resistance of 3-unit Y-TZP frameworks was significantly
reduced after veneering with a feldspar-based ceramic or a glass ceramic or after heat
treatment in a way similar to veneering (Sundh et al, 2005). Similarly, the fracture resistance
of an Mg-PSZ (Denzir-M®) ceramic was significantly reduced by heat treatment or veneering,
while that of a Y-TZP (Vita CerecInLab®) ceramic considerably increased after veneering
(Sundh and Sjogren, 2006). Although the reported results differ, the mechanical properties of
zirconia ceramics are clearly affected by surface treatments (Zhang et al, 2004; Guazzato et
al, 2005b; Sundh et al, 2005; Sundh and Sjogren, 2006). In the present study, all Y-TZP
frameworks were veneered with a compatible feldspar-ceramic, and then tested to fracture.
There has been no control group of as-sintered (without veneering) frameworks. Therefore,
A goal of prosthetic dentistry is to preserve the remaining tooth structure even if the
biologi l and biophysical failures include fracture of the metal and/or
eramics, abutment fracture, and loss of retention (Strub et al, 1988).
the influence of the veneering process on fracture resistance of the tested materials could not
be evaluated.
7.2.5 Fracture patterns
restoration fails. Failures of a bridge can occur because of technical, biophysical and
cal problems. Technica
c
For most of the specimens in the present study, the fractures were located at the loading point
and through one or both of the connectors. One specimen of the non-aged DC-Zirkon®
exhibited a tooth fracture; while 3 specimens, 1 of the aged DC-Zirkon® group and 2 of the
aged Vita CerecInLab- group, showed fractures of the self-curing resin (Technovit 4000,
Zwick, Ulm, Germany) surrounding the premolar abutments. In general, the fracture patterns
observed were similar to those reported in previous in vitro studies of 3-unit partially-
stabilized zirconia FPDs (Filser et al, 2001; Tinschert et al, 2001a; Sundh et al, 2005). It was
difficult to assess the exact origin of the fractures occurred; whether they start at the loading
point or at the connectors. The fractures, however, were perpendicular to the mesial-distal
axis of the frameworks in a smooth curve between the loading point and the gingival side of
the connector. The failure mode is in agreement with findings of previous studies showing
71
DISCUSSION
that the exclusive mode of failure in vitro and in vivo for all-ceramic FPDs was a fracture of
the connectors, and that the gingival side of the connectors can be the area where high tensile
stresses are located (Campbell and Sozio, 1988; Kelly et al, 1995; Filser et al, 2001;
Tinschert et al, 2001a; Vult von Steyern et al, 2001; Sundh et al, 2005).
To ensure long-term success of PFMs, the minimal critical dimensions recommended for the
connectors are 2.5 x 2.5 mm (Miller, 1977; McLean, 1982). For all-ceramic FPDs, due to
their primary mode of failure and their brittleness, the required connector dimensions are
larger. This may be a major contributing factor in restricting the versatility of their use.
Therefore, appropriate diagnosis, patient selection and conception of the requirements, such
as anatomical limitations, hygenic reasons and esthetic expectations of proper ceramic
framework design, are crucial for the success of these restorations (Raigrodski, 2004a). In the
present study, connector dimensions of 3x3 mm were used. Several in vitro studies,
investigating the fracture resistance of Y-TZP based all-ceramic FPDs with connector
dimensions of 3x3 mm, have yielded good results (Vult von Steyern, 2005; Vult von Steyern et
al, 2005) In a clinical study, Y-TZP all-ceramic FPDs with connector dimensions of 4x4 mm
showed a success rate of 100% after an observation period of 2 years (Vult von Steyern et al,
005). No clinical studies are available on the clinical success of Y-TZP all-ceramic FPDs
with connector dimensions of 3x3 mm. Therefore the long-term success of these restorations
ommending them for clinical application.
2
should be evaluated before rec
72
CONCLUSIONS
8 Conclusions
Within the limitations of this in vitro study, the following conclusions may be drawn:
1. All tested Y-TZP all-ceramic FPDs have the potential to withstand physiological occlusal
forces applied in the posterior region, and appear therefore to be a viable alternative to replace
conventional posterior PFMs.
2. Critical issues such as connector shape and size, veneering ceramics, aging behavior and
long-term clinical performance need to be further assessed before recommending Y-TZP all-
ceramic FPDs for daily practice.
73
SUMMARY
9 Summary
Objective: The purpose of this in vitro study was to evaluate the fracture resistance of
ifferent zirconia three-unit posterior all-ceramic fixed partial dentures (FPDs) before and
riodontal membrane, in order to simulate the physiological
butment mobility, and fixed into models representing a 3-unit FPD. After impression taking
® Cubes,Vita Zahnfabrik, D-Bad Säckingen) were
bricated for the 3-unit FPDs (3 groups, 16 FPDs each). All frameworks were veneered with
e e-max Ceram® ceramic (Ivoclar, Vivadent, FL-Schaan). All FPDs were subsequentally
emented with glasionomer-cement (Ketac Cem®, GC Europe, B-Leuven). One-half of the
pecimens (n=24) were artificially aged in the chewing simulator (1,2 million cycles, 49 N).
ll of the samples were then loaded until fracture occurred using a universal testing machine
wick Z010/TN2S®, Zwick, Ulm, Germany).
esults: All specimens, subjected to artificial aging, survived with no failures. The median
acture resistance values (min, max) before artificial aging were for DCS: 1683 N
278,1993); Procera: 1522 N(1105, 1800); Vita: 1702 N (1472, 1946) - and after aging for
CS:1618 N (1175, 1804); Procera:1256 N (1044, 1783); and Vita: 1556 N (1394, 1854). The
ffect of artificial aging was not statistically significant between the test groups. Similarly, no
ignificant differences were found for the fracture resistance comparisons between different
roups before artificial aging. After artificial aging, Procera showed significantly smaller
acture resistance than Vita (Wilcoxon test: p=0.015) and DCS (Wilcoxon test: p=0.038).
onclusions: All tested restorations have the potential to withstand occlusal forces applied in
e posterior region and may thus represent viable alternatives for use as posterior all-ceramic
storations. Other issues such as connector size and shape, ceramic veneering materials and
ethods, aging behavior, and long-term clinical performance need to be further assessed
efore recommending such restorations for daily practice.
d
after fatigue loading in the chewing machine.
Material and methods: 96 teeth (48 mandibular premolars and 48 molars) were prepared,
covered with an artificial pe
a
and the fabrication of master models, 48 frameworks from 3 different zirconiumdioxide
materials (Procera Zirconia®, Nobel-Biocare AB, S-Göteborg; DC-Zirkon®, DCS Dental AG,
CH-Allschwill; Vita In-Ceram YZ
fa
th
c
s
A
(Z
R
fr
(1
D
e
s
g
fr
C
th
re
m
b
74
ZUSAMMENFASSUNG
10 Zusammenfassung
evaluieren.
wick Z010/TN2S®, Zwick, Ulm,
on
assen, Alterung und langzeitiges
linisches Verhalten sollten noch weitergeprüft werden, bevor diese Restorationen als
lltägliches Verfahren empfohlen werden können.
Zielsetzung: Das Ziel der vorliegenden in vitro Studie war es die Bruchfestigkeit
von verschiedenen dreigliedrigen vollkeramischen Zirkondioxid Seitenzahnbrücken vor und
nach künstlicher Alterung im Kausimulator zu
Material und Methode: Es wurde 96 Zähne (48 Unterkieferprämolaren und 48 Molaren)
beschliffen und mit einem künstlichen parodontalen Ligament, welches die physiologische
Mobilität des Zahnes simuliert, in Modelle, entsprechend einer dreigliedrigen Brücke, fixiert.
Nach Abformung und Arbeitsmodellherstellung wurden 48 Gerüste für dreigliedrige
Seitzenzahnbrücken (3 Gruppen à 16 Brücken) aus drei verschiedenen Zirkoniumdioxid
Materialien angefertigt: Procera Zirconia® (Nobel-Biocare AB, S-Göteborg), DC-Zirkon®
(DCS Dental AG, CH-Allschwil), Vita In-Ceram YZ® Cubes (Vita Zahnfabrik, D-Bad
Säckingen).Alle Gerüste wurden mit e-max Ceram® Keramik (Ivoclar, Vivadent, FL-Schaan)
verblendet. Die Seitenzahnbrücken wurden mit Glasionomer-Zement (Ketac Cem®, GC
Europe, B-Leuven) zementiert. Jeweils eine Hälfte (n=24) der Prüfkörper wurde im
Kausimulator künstlich gealtert (1,2 Millionen Zyklen, F= 49 N). Alle Prüfkörper wurden bis
zum Bruch belastet (Universal-Prüfmaschine: Z
Deutschland).
Ergebnisse: Alle Prüfkörper hielten der Kausimulation stand. Die mediane Bruchfestigkeit
(Min, Max) betrug ohne Kausimulation für Procera 1256 N (1105, 1800), DCS 1522 N (1278,
1993), Vita 1618 N (1472, 1946), und nach künstlicher Alterung für Procera 1256 N (1044,
1783), DCS 1618 N (1175,1804), Vita 1556 N (1394, 1854). Innerhalb einer Gruppe, der
Einfluß der künstlichen Alterung war nicht statistisch signifikant. Es gaben ebenfalls keine
signifikante Unterschiede für die Bruchfestigkeit zwischen den Gruppen ohne Alterung. Nach
künstlicher Alterung, Procera zeigte signifikant niedrigere Bruchfestigkeit als DCS (Wilcox
test: p=0.038) und Vita (Wilcoxon test: p=0.015).
Schlussfolgerung: Alle getesteten Restorationen hielten den im Seitenzahnbereich
Belastungskräften stand und könnten als interessante Alternative zur konventionellen
metallkeramischen Versorgungen im Seitenzahnbereich in Betracht gezogen werden.
Zusätzliche Parameter, wie Verbindergestaltung, Verblendm
k
a
75
APPENDIX
11 Appendix
11.1 Fracture resistance values of the Procera group
Without aging (Tab. 11.1)
Sample # 1 2 3 4 5 6 7 8
F (N) 1714. 48 1105.26 1254.51 1765.86 1285.28 1559.14 1799.90 1484.78
Table 11.1
With aging (Tab. 11.2)
Sample # 1 2 3 4 5 6 7 8
F (N) 1252.64 1783.25 1503.52 1094.96 1043.84 1258.92 1295.77 1142.68
Table 11.2
11.2 Fracture resistance values of the DCS group
Without aging (Tab. 11.3)
Sample # 1 2 3 4 5 6 7 8
F (N) 1421.87 1612.77 1511.88 1277.84 1993.08 1859.20 1754.18 1840.76
Table 11.3
76
APPENDIX
With aging (Tab. 11.4)
Samp 7 8 le # 1 2 3 4 5 6
F(N) 1545.00 1454.73 1174.66 1713.17 1714.10 1637.74 1804.15 1598.49
Table 11.4
ctu resis nce v es of e Vita roup
ithout aging (Tab. 11.5)
3 4 5 6 7 8
11.3 Fra re ta alu th g
W
Sample # 1 2
F(N) 1755.66 1615.08 1471.90 1673.49 1691.8 1945.80 1837.16 1712.56
Table 11.5
ith aging (Tab. 11.6)
mple # 1 2 3 4 5 6 7 8
W
Sa
F(N) 1600.5 1571.59 1853.76 1394.01 1443.79 1541.18 1852.25 1490.34
Table 11.6
77
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98
CURRICULUM VITAE
13 Curriculum vitae
Date of Birth 16. October 1977
Place of Birth Aigio, Achaia, Griechenland
umana-Stamouli
nality Greek
1988-1989 Secondary School, Patras, Greece
Wuppertal, Germany
ty 1995-1996 Chemical engineering, Metsoveio
ment 2001-2002 Assistant in private dental office,
2003-2006 Posttgraduate student, Department of
Prosthodontics, Albert Ludwigs
University, Freiburg, Germany
Parents Vasileios Stamoulis
Maria Ro
Marital Status Single
Natio
Education 1983-1988 Secondary School, Aigio, Greece
1989-1991 Gymnasium, Patras, Greece
1991-1992 Gymnasium,
1992-1995 Lyzeum, Wuppertal, Germany
Universi
Polytechnik Institute, Athens, Greece
1996-2001 Dental School, University of
Athens, Greece
Employ
Athens, Greece
2002-2003 Assistant in private dental office,
London, England
99
ACKNOWLEDGEMENTS
14 Acknowledgements
J. R. Strub, Chair, Department
f Prosthodontics, Albert Ludwigs University, Freiburg, Germany for offering me the
arry o resea h und
GmbH in der Helmholtz-Gemeinschaft,
e manuscript.
r. W. Att, The Weintraub Center for Reconstructive Biotechnology, UCLA School of
os ngele Calif rnia, U
throughout the study.
Witko
Germany, for his help and technical
r. M. Tomic, Department of Prosthodontics, Albert Ludwigs University, Freiburg, Germany
igs
ity, Fre erman
he dental technicians of Nobel Biocare, Goetheborg, Sweden; Mr Ahlmann, Chief, Dental
kheim, Germ y and ba h,
he dental laboratory Woerner Zahntechnik, Freiburg, Germany for the veneering of the test
samples.
would like to add a special thank to my parents for their continuous moral and financial
support throughout my studies.
I would like to express my most sincere gratitude to Prof. Dr.
o
opportunity to c ut rc er his supervision.
I would also like to thank:
Prof. Dr. J. Hausselt, Forschungszentrum Karlsruhe
Materials Research Department, for the review of th
D
Dentistry, L A s, o SA, for his continuous advise, help and support
ZTM S. wski, Department of Prosthodontics, Albert Ludwigs University, Freiburg,
support throughout the experiment.
D
for the conception of the idea of the present study.
Dr. T. Gerds, Institute of Medical Biometry and Medical Informatics, Albert Ludw
Univers iburg, G y, for the statistical analysis of the data.
T
Laboratory, Kel an Mr G. Lom rdi, Chief, Dental laboratory, Zuric
Switzerland, for the fabrication of the frameworks of the test samples.
T
I
I would also like to thank Jörg for his patience, support and understanding even under
difficult circumstances.
100
ACKNOWLEDGEMENTS
101
