Survival rate and fracture resistance of zirconium dioxide implants ...

$: Survival rate and fracture resistance of zirconium dioxide implants ...$
Aus der Universitätsklinik für Zahn-, Mund-, und Kieferheilkunde der

Albert-Ludwigs Universität Freiburg i. Br.

Abteilung für Zahnärztliche Prothetik

(Ärztl. Direktor: Prof. Dr. J. R. Strub)

Survival rate and fracture resistance of zirconium

dioxide implants after exposure to the artificial mouth:

An in-vitro study

INAUGURAL- DISSERTATION zur Erlangung des

Zahnmedizinischen Doktorgrades

Der Medizinischen Fakultät der Albert- Ludwigs- Universität

Freiburg i. Br.

Vorgelegt 2006

Von Marina Andreiotelli

Geboren in Athen, Griechenland

Dekan: Prof. Dr. C. Peters

1. Gutachter: Prof. Dr. R. J. Kohal

2. Gutachter: PD Dr. T. Auschill

Promotionsjahr: 2006

Table of contents

1. INTRODUCTION.......................................................................................... 1

2. LITERATURE REVIEW ............................................................................ 4

2.1 HISTORY OF IMPLANTS AND THEIR USE IN DENTISTRY............................................ 4

2.2 OSSEOINTEGRATION AND CRITERIA OF SUCCESS .................................................... 6

2.3 IMPLANT MATERIALS............................................................................................... 7

2.4 SURVIVAL RATES OF SINGLE-TOOTH IMPLANT REPLACEMENTS............................. 9

2.5 PURE TITANIUM AND ITS ALLOYS........................................................................... 11

2.5.1 Mechanical properties of cp titanium and alloys ............................................. 14

2.5.2 Corrosive behavior and biocompatibility of pure titanium and its alloys........ 15

2.6 THE USE OF ZIRCONIUM DIOXIDE IN DENTISTRY ................................................... 18

2.6.1 Zirconium and the mineral Zircon ................................................................... 20

2.6.2 The mineral zirconium dioxide........................................................................ 22

2.6.3 Zirconium dioxide as a high-performance ceramic biomaterial ...................... 23

2.6.4 Mechanical properties and ageing of Y-TZP................................................... 25

2.6.5 Biocompatibility of Y-TZP.............................................................................. 28

3. AIM OF THE STUDY ................................................................................. 30

4. MATERIALS AND METHODS ........................................................... 31

4.1 OUTLINE OF THE STUDY ......................................................................................... 31

4.2 MATERIALS USED FOR THE FABRICATION OF THE IMPLANTS ............................... 33

4.2.1 Y–TZP BIO-HIP Sigma® Implants (Incermed, CH-Lausanne)....................... 34

4.2.2 Y-TZP-A BIO-HIP® Implants (Metoxit AG, CH-Thayngen) ......................... 35

4.3 IMPLANT HEAD PREPARATION ............................................................................... 38

4.4 FINAL IMPRESSIONS AND FABRICATION OF THE CROWNS ..................................... 39

4.5 LUTING PROCEDURES ............................................................................................. 40

4.6 FABRICATION OF THE MASTER MODELS FOR THE ARTIFICIAL MOUTH ................ 41

4.7 EXPERIMENTS ......................................................................................................... 43

4.7.1 Exposure of the test samples to the artificial mouth ........................................ 43

4.7.2 Fracture strength test (Fig. 21)......................................................................... 47

4.7.3 Statistical analysis of data................................................................................ 48

5. RESULTS............................................................................................................. 49

5.1 DYNAMIC LOADING IN THE ARTIFICIAL ORAL ENVIRONMENT.............................. 49

5.1.1 Survival rate of the implants ............................................................................ 49

5.2 FRACTURE STRENGTH TEST ................................................................................... 51

5.2.1 Fracture strength of the implants ..................................................................... 51

5.2.2 Statistical evaluation of the data ...................................................................... 55

5.2.3 Fracture patterns of the samples ...................................................................... 57

6. DISCUSSION..................................................................................................... 59

6.1 ZIRCONIUM DIOXIDE ONE-PIECE IMPLANTS AS MATERIAL OF INTEREST ............. 59

6.2 OVERVIEW OF THE RESULTS .................................................................................. 62

6.3 OCCLUSAL FORCES IN THE ANTERIOR REGION – CLINICAL RELEVANCE WITH THE

FRACTURE STRENGTH VALUES OF CERAMIC MATERIALS ........................................... 63

6.4 THE CLINICAL RELEVANCE OF FRACTURE STRENGTH TESTS................................ 65

6.5 CLINICAL RELEVANCE AND INFLUENCE OF THE ARTIFICIAL MOUTH ON THE...... 66

SURVIVAL RATE AND FRACTURE STRENGTH OF ZRO2 IMPLANTS ............................... 66

6.6 SURFACE AND HEAT TREATMENTS OF ZIRCONIA DEVICES (ABUTMENTS, CROWNS)

....................................................................................................................................... 70

6.7 INFLUENCE OF THE SURFACE TOPOGRAPHY AND THE DESIGN OF THE IMPLANTS

ON THE FRACTURE STRENGTH VALUES........................................................................ 72

6.8 PATTERNS OF FRACTURE OF THE SAMPLES ........................................................... 73

7. CONCLUSIONS.............................................................................................. 75

8. SUMMARY......................................................................................................... 76

9. ZUSAMMENFASSUNG............................................................................. 77

10. REFERENCES............................................................................................... 78

11. CURRICULUM VITAE......................................................................... 100

12. ACKNOWLEDGEMENTS.................................................................. 101

To my family Στην οικογένειά μου

Introduction 1

1. Introduction

Over the last few years, the success of osseointegrated dental implants has revolutionized

dentistry. The rehabilitation with the use of implants has become a well-accepted

treatment modality, as the ability to permanently replace missing teeth in edentulous and

partially edentulous situations as well as in situations were only a single tooth is missing,

with a function and appearance close to that of the natural dentition, has never been

greater (van Steenberghe 1989; Adell et al. 1990; Henry et al. 1996).

With more than three decades of evidence to support the clinical use of osseointegrated

dental implants made of pure titanium, it is possible to confidently confirm that these

implants are predictable and provide patients with long-term functional tooth replacement

(Albrektsson et al. 1988; Adell et al. 1990; Binon 2000). This is a remarkable

accomplishment, considering the many challenges and stresses that the oral environment

and forces of mastication present for dental implants.

Replacing single missing teeth, especially in the anterior region, has always been a

challenge for dentists. With increasing patient demands, removable partial dentures have

become less acceptable and many patients now oppose the preparation of intact teeth for

the fabrication of a fixed partial denture. Other treatment options, such as resin-bonded

restorations, orthodontics, and tooth transplantation have been created for replacing

single missing teeth. However, none of these alternatives have proved to be ideal, and

thus there was a great interest in replacing single missing teeth with implant-supported

crowns (Ekfeldt et al. 1994; Andersson et al. 1998).

Simultaneously the expectations regarding esthetics in dentistry are growing and research

in the field of all-ceramic materials for restoration of the natural dentition and dental

implants has been intensified. During the past two decades numerous types of ceramics

(i.e. IPS-Empress, Empress 2, In-Ceram Alumina, In-Ceram Spinell and In-Ceram

Zirconia, aluminum oxide, zirconium dioxide ceramic) (Kappert and Krah 2001;

Tinschert et al. 2001b) and novel processing methods have been introduced for the

fabrication of crowns, bridges, inlays, onlays, veneers as well as for the reconstruction of

Introduction

2

dental implants. The introduction of computer-aided design and manufacturing (CAD-

CAM) has become an increasingly interesting alternative to manual, casting, or pressing

techniques (Tinschert et al. 2004; Witkowski 2005).

Although the crown that covers the abutment of an implant may be esthetically optimal,

the possibility exists that the grayish of the titanium implant shines through the thin

periimplant mucosa, thus impairing the entire esthetic result (Wohlwend et al. 1996;

Heydecke et al. 1999). Furthermore, due to soft peri-implant tissue recessions, the

implant head might become visible over time.

There are reports that metals including titanium are able to induce non-specific

immunomodulation and autoimmunity (Stejskal and Stejskal 1999). With highly sensitive

immunologic in vitro tests sensitization to titanium could be observed in some cases

(Lalor et al. 1991; Valentine-Thon and Schiwara 2003). The clinical relevance of the

above findings is not clear so far.

These are some reasons why nowadays a serious effort is made to create implants that are

more “patient friendly”, maintaining at the same time the characteristics giving them high

success rates (Yildirim et al. 2000). A ceramic implant could solve the possible

mentioned “health” and esthetic problems with titanium implants and could be a viable

alternative, especially in maxillary anterior regions (Kohal and Klaus 2004).

However, a major problem with all-ceramic restorations is the low fracture resistance.

This disadvantage made it difficult to apply all-ceramics for the fabrication of dental

implants. One such implant material, aluminum oxide, was used, for example, with the

Tübingen Implant (Frialit I) (Schulte and Heimke 1976), but because of its insufficient

physical properties it was withdrawn from the market.

Recently, another ceramic material with potential for future use as dental implant

material was introduced. Zirconium dioxide (ZrO2) as metal substitute possesses good

physical characteristics, like a high flexural strength (900-1200 MPa), Vickers hardness

(1200), and Weibull modulus (10-12) (Marx 1993; Piconi and Maccauro 1999). Because

of its stability, its good mechanical properties and high biocompatibility this material

could possibly open a new way to implant dentistry. The application of ZrO2 for the

fabrication of dental implants has been previously suggested (Albrektsson et al. 1985).

Introduction

3

To date, data regarding the use of zirconium oxide in dental implantology only exist in

animal and laboratory trials. Long-term data regarding the use under clinical conditions

do, however, not exist.

Literature Review 4

2. Literature Review

2.1 History of implants and their use in dentistry For centuries, people have attempted to replace missing teeth using transplantation and/or

implantation. The origins of dental implants began as early as the ancient Greeks,

Etruscans and Egyptians. These civilizations employed different implantation materials

ranging from jade and bone to metal. Some of the designs used have evolved into the

modern implants we see today (Marziani 1954).

In the 20th century, Greenfield introduced in 1906 for the first time a hollow basket-

shaped implant made of an iridium-platinum alloy. Greenfield’s implant model, with its

rather unusual design, could be considered a prototype of the hollow cylinder implants

still used today. His seven-year research work was first acknowledged by the Academy of

Stomatology in Philadelphia (Greenfield 1991).

In the early 1930s more emphasis was placed on the tissue tolerance as well as the bone

reaction towards metal implants. The introduction of stainless steel and the development

of a cobalt-chromium-molybdenum alloy (Vitallium) gave new impulses to implant

surgery. Strock succeeded in anchoring a Vitallium screw within bone and immediately

mounting a porcelain crown to the implant. Vitallium was considered to be inert,

compatible with the living tissue, and resistant to body fluids. Strock for the first time

achieved a long term implant survival for over fifteen years. He stated that good

occlusion was critical to prevent traumatization of the implant and avoid unwanted

resorption processes (Strock 1939).

At the same time that Strock made his first attempts, Müller followed a different path. In

1937, he placed the first subperiostal implant, made of an iridium-platinum alloy and left

four abutments protruding into the oral cavity. At first it seemed that his concept of

leaving the inner bone structures undisturbed and placing the foreign body between the

bone and periosteum corresponded more to the physiologic conditions than did

endosseous insertion techniques. However, later developments showed that his concept

could not live up to these expectations (Müller 1937).

Literature Review 5

In the 1950s and 1960s, numerous implantologists developed subperiostal implant

procedures, among them Marziani, who in 1955, returned to an one-step procedure by

placing a perforated deformable tantalum plate (Marziani 1954). Later combinations of

subperiostal and endosseous implant systems such as the mandibular staple bone blade by

Small and their modifications could not achieve universal acceptance (Small 1975).

During the past 30 years, Per-Ingvar Brånemark and Leonard I. Linkow have markedly

influenced the development of implant surgery. Linkow could be considered as one of the

most creative implantologists. He developed a blade-type perforated endosseous implant

and he was the first that aimed to increase the contact area between the implant surface

and the peri-implant bone and to adapt implants to the respective anatomic conditions

with minimal surgical burden on the patient (Linkow 1970).

In 1952, Brånemark, the father and mentor of modern implant surgery, developed a

threaded implant design made of pure titanium that increased the popularity of implants

to new levels. Discovering almost by chance the high compatibility and strong anchorage

of titanium in the bone marrow of rabbit fibula, he and his co-workers were the first to

introduce the term osseointegration, defined as the tenacious affinity between living bone

and the titanium oxides (Brånemark 1983). He continued his basic research and clinical

examinations, studying every aspect of implant design, including biological, mechanical,

physiological, and functional phenomena (Brånemark et al. 1977). His research ended in

a two-stage implant system for oral endosseous implantations, which was not marketed

until 17 years of extensive clinical testing and study has been completed.

In the later 1970s and early 1980s the research group of Schroeder et al. examined also

the functional ankylosis of titanium implants in the jawbone. In an animal experiment

they reported on a non-submerged approach to dental implant placement and described

the soft tissue attachment/contact to the transgingival portion of a new introduced hollow

cylinder titanium implant with a special surface texture (titanium plasma sprayed)

(Schroeder et al. 1976; Schroeder et al. 1978; Schroeder et al. 1981).

Since 1970, numerous new implant materials and designs followed, including the use of

polymers, porcelain, high-density aluminum oxide (alumina), bioactive glass (Bioglass)

and carbon.

Literature Review 6

In 1970, Hodosch et al. described a root-shaped implant made of polymethylmethacrylate

(Hodosch et al. 1970). In 1976, Schulte and Heimke introduced the Tübingen immediate

implant, which could be used for immediate restoration of an extracted or lost tooth, and

was made of an aluminum oxide ceramic material (Schulte and Heimke 1976; Schulte

1984a, b, c).

Nowadays, the most frequently used implant material is titanium. As a result of

Brånemark’s extensive studies, titanium has become the gold standard in implant

dentistry. However the great revolution in the field of ceramic materials, with the use of

zirconium dioxide could open a new, challenging way in implantology.

2.2 Osseointegration and criteria of success

Implantation is defined as the insertion of any object or material, such as an alloplastic

substance or other tissue, either partially or completely, into the body for therapeutic,

diagnostic, prosthetic, or experimental purposes (Marziani 1954). Dental implants, which

are also implantation materials, have a form and shape that reminds of natural teeth as

they are used to replace teeth that are missing in the natural dentition.

Brånemark and his colleagues introduced the term osseointegration in 1977 to describe

direct bone anchorage of dental titanium screw implants that can withstand long-term

functional loading (Brånemark et al. 1977). It is described further as the direct adaptation

of bone to implants without any other intermediate interstitial tissue, and it is similar to a

tooth ankylosis were no periodontal ligament exists. By definition, osseointegration

demands the absence of a fibrous layer, and implies that the biological response of the

bone is not one of inertness toward a foreign material but rather one of integration of the

material with the bone as if it is a part of the body (Brånemark 1983; Wataha 1996).

To be considered successful, an osseointegrated oral implant has to meet certain criteria

in terms of function (ability to chew), tissue physiology (presence and maintenance of

osseointegration, absence of pain and other pathological processes) and user satisfaction

(esthetics and absence of discomfort) (Esposito et al. 1998). The often misquoted term of

implant survival relates only to the devices remaining in the jaws of the patient, while the

quality of the survival as well as the function of the implant is irrelevant. If an implant is

Literature Review 7

still in the jaw but is not tested for osseointegration, as it happens in the case of implants

lost to follow-up and of “sleeping” implants, this implant may be categorized as surviving

but not successful (Albrektsson and Sennerby 1991).

By today’s standards, the presence of a fibrous layer between bone and a dental implant

indicates failure of the implant. For an implant system to be successful several

requirements have been set up: 1) immobility in any direction, 2) the average

radiographic marginal bone loss should be less than 1.5 mm during the first year of

function and less than 0.2 mm annually following thereafter and 3) the radiograph should

not demonstrate any evidence of periimplant radiolucency. Finally, the individual implant

performance must be characterized by absence of signs and symptoms such as pain,

infection, paresthesia, or violation of the mandibular canal. Success rates of 85% or more

at the end of a five-year observation period and 80% at the end of a ten-year period are

the minimum criteria for implant success (Albrektsson et al. 1986; Shulman et al. 1986;

Albrektsson and Lekholm 1989; Albrektsson and Sennerby 1991). Smith and Zarb

modified Albrektsson’s criteria by stating that the patient’s and dentist’s satisfaction with

the implant prosthesis should be the primary consideration and that esthetic requirements

should also be met (Schmitt and Zarb 1993).

2.3 Implant Materials A wide variety of materials have been used in dental and maxillofacial implants. They

can be classified as metals and alloys, polymer-based materials, ceramics, glasses and

carbons (Table 1) (Lemons 1990).

Implant material Common name or abbreviation

Metals

Titanium Cp Ti (commercially pure titanium)

Titanium Alloy Ti 6Al 4V

Stainless Steel SS, 316 L SS

Cobalt Chromium Alloy Vitallium, Co-Cr-Mo

Literature Review 8

Gold Alloys

Tantalum Ta

Ceramics

Alumina Al2O3, amorphous or single crystal

sapphire (Bioceram®, Kyocera, J-Kyoto)

Hydroxyapatite HA, Ca10 (PO4) 10(OH) 2

Beta-Tricalcium phosphate β-TCP, Ca3 (PO4) 2

Carbon C, vitrous, low temperature isotropic

(LTI), ultra-low temperature isotropic (ULTI)

Carbon-Silicon C-Si

Bioglass SiO2 / CaO / Na2O / P2O5

Polymers

Polymethylmethacrylate PMMA

Polytetrafluoroethylene PTFE

Polyethylene PE

Polysulfone PSF

Polyurethane PU

Table 1 Materials that have been used for endosseous implants (Williams 1981; Lemons

1990; Wataha 1996)

As with all implanted devices the requirements of the materials used for their

construction vary but can be broadly classified under the headings of biocompatibility,

biofunctionality and availability. Biocompatibility refers to the interactions between

materials and the tissues of the body and is one of the most important factors involved

with the material selection (Williams 1981). Biofunctionality is concerned with those

mechanical and physical properties that enable the implanted device to perform its

function under the stresses imposed on dental implants in the oral cavity. Availability

refers to the techniques of fabrication and sterilization of the implants (Williams 1981).

In the above mentioned categories of implant materials, the low mechanical strength of

most of them (ceramics, polymers, carbon), results in a greater susceptibility for

Literature Review 9

mechanical fracture during function. Simultaneously, the inferior to today’s standards

biocompatibility of some of them (i.e. stainless steel, cobalt-chromium alloys) precludes

their use in implant dentistry, as they promote the formation of a fibrous interface with

bone (Wataha 1996).

Nowadays, the most popular implant material in use is commercially pure titanium and

some of its alloys, mainly because of its several favorable physical and mechanical

properties and its biocompatibility.

2.4 Survival rates of single-tooth implant replacements

The longevity of dental titanium implants is an important concern. Several studies have

been published investigating the long-term prognosis of titanium dental implants.

However, to date, long-term documentation referring to implant treatment in partially

edentulous cases, including loss of single teeth, is still not as extensive as for the

completely edentulous jaw (Lindh et al. 1998).

It must be noted that most publications, which have reported on survival or success rates

of single implants and attached prosthetic components, are based on studies of

Brånemark implants (Schmitt and Zarb 1993; Ekfeldt et al. 1994; Engquist et al. 1995;

Avivi-Arber and Zarb 1996; Henry et al. 1996; Malevez et al. 1996; Parein et al. 1997;

Andersson et al. 1998; Scheller et al. 1998; Polizzi et al. 1999; Scholander 1999; Bianco

et al. 2000; Naert et al. 2000; Haas et al. 2002).

Results for other implant systems (Frialit-2, ITI, Astra Tech Implants etc.) have also

documented the successful use of implant systems for single-tooth replacement (Gomez-

Roman et al. 1997; Levine et al. 1999; Priest 1999; Palmer et al. 2000; Norton 2001;

Krennmair et al. 2002).

In Table 2 some of the longitudinal studies are listed concerning single-tooth

replacements through oral implants.

Literature Review 10

Author (year) Study

Design

Observation

period (years)

Type of

implants

Number of

implants

Survival Rate*/

Success Rate**

(%)

Ekfeldt et al.

(1994)

Retrospective 1-4.5 Brånemark 93 98 *

Engquist et al.

(1995)

Retrospective 1-5 Brånemark 82 97.6 *

Avivi-Arber et al.

(1996)

Prospective 1-8 Brånemark 49 98 *

Henry et al.

(1996)

Prospective 5 Brånemark 107 Maxillae: 96.6 **

Mandible: 100 **

Malevez et al.

(1996)

Retrospective >5.1 Brånemark 84 97.6 *

Gomez-Roman

et al. (1997)

Prospective 1-5 Frialit-2 290 97.6 *

Parein et al.

(1997)

Retrospective 6 Brånemark 56 89 *

Andersson et al.

(1998)

Prospective 5 Brånemark 65 98.5 **

Scheller et al.

(1998)

Prospective 1-5 Brånemark 99 Maxillae: 97.4 **

Mandible: 83.3 **

Polizzi et al.

(1999)

Retrospective 5 Brånemark 30 93.3 **

Priest (1999) Prospective 10 3i (100), Nobel

Biocare(12)

Steri-Oss(2),

Impla-Med(1),

Friatec (1)

116 97.4 *

Scholander

(1999)

Retrospective 1-9 Brånemark 259 98.3 **

Bianco et al.

(2000)

Retrospective 8 Brånemark 252 Maxillae: 96.4 *

Mandible: 94.7 *


Naert et al.

(2000)

Prospective 1-11 Brånemark 270 93 **

Palmer et al.

(2000)

Prospective 5 Astra Tech 14 100 *

Norton (2001) Prospective 1-7 Astra Tech 27 95.6**

Haas et al.

(2002)

Prospective 0.5-10 Brånemark 76 93 *

Krennmair et al.

(2002)

Retrospective 0-7 Frialit-2 146 97.3 *

Table 2 Survival/success rates of single-tooth implant replacements (* = survival rate, ** = success

rate)

These studies illustrate that the survival rate of single-tooth implant replacements ranges

between 93% and 100% for observation periods of up to eleven years, revealing very

good clinical success. A meta-analysis by Lindh and colleagues demonstrated a 97.5%

cumulative survival rate after 6 to 7 years, among 570 single implants from nine studies

with a 1- to 8-year loading time (Lindh et al. 1998). Another systematic review by

Creugers et al. also based on nine studies showed a survival rate of 97±1% for 459 single

teeth implants after 4 years observation period (Creugers et al. 2000).

Other studies with shorter observation periods between one to three years have evaluated

single titanium fixtures of various systems and reported good results as well (Jemt and

Pettersson 1993; Becker and Becker 1995; Levine et al. 1999).

2.5 Pure titanium and its alloys

As a result of Brånemark’s studies, titanium has become the most commonly used

implant material in dentistry.

The element was discovered by Wilheim Gregor, a clergyman, who found the metal in

“black magnetic sand” in Cornwall in 1791. Three years later, Klaproth found a rutile that

was the oxide of a new metal he named titanium, after the Greek Titans. He recognized


that this metal was identical to the material Gregor had discovered (Weber et al. 1992;

McCracken 1999).

Titanium exists in nature as a pure element with an atomic number of 22 in the periodic

table and an atomic weight of 47.9. It is the ninth most abundant element and the fourth

most abundant structural metallic element in the earth’s crust, following aluminum, iron and magnesium. About 0.6% of the earth’s crust is composed of titanium, and it is a

million times more abundant than gold. Large reserves of this metal are found in Canada,

Australia, and the United States. Most titanium used today comes from mines located in

Australia. Of the total amount of titanium mined, only 5% to 10% is used in its metal

form (Parr et al. 1985).

The metal exists as rutile (TiO2) (Fig. 1) or ilmenite (FeTiO3) (Fig. 2) and requires

specific extraction methods to be recovered in its elemental state. It is produced by

heating titanium ore in the presence of carbon and chlorine and then reducing the

resultant TiCl4 with molten sodium to produce a titanium sponge. This sponge is then

fused under vacuum or in an argon atmosphere into ingots composed of the familiar

metal (Parr et al. 1985; Lautenschlager and Monaghan 1993).

Fig. 1 The mineral rutile (TiO2) Fig. 2 The mineral ilmenite (FeTiO3)

At temperatures up to 882oC, pure titanium exists as a hexagonal close-packed atomic

structure (alpha phase). Above that temperature, the structure is body-centered cubic

(beta phase). The metal melts at 1,665oC (McCracken 1999).

Titanium alloys of interest to dentistry exist in three forms: alpha, beta and alpha-beta.

These types originate when pure titanium is heated, mixed with elements such as


aluminum and vanadium in certain concentrations, and then cooled. Titanium can be

alloyed with a wide variety of elements to alter its properties, mainly for the purposes of

improving strength, high temperature performance, creep resistance, response to ageing

heat treatments and formability (Lautenschlager and Monaghan 1993). The elements

oxygen, aluminum, carbon, and nitrogen stabilize the alpha phase of titanium because of

their increased solubility in the hexagonal close-packed structure. Oxygen occupies

interstitial sites. Aluminum also serves to increase the strength and decrease the weight of

the alloy. Elements that stabilize the beta phase include magnesium, chromium, iron, and

vanadium (Wataha 1996). These beta-phase stabilizers are used to minimize the

formation of TiAl3 to approximately 6% or less to decrease the alloys susceptibility to

corrosion (McCracken 1999).

The alloys most commonly used for dental implants are of the alpha-beta variation. Of

these, the most common contains 6% aluminum and 4% vanadium (Ti 6Al 4V). After

heat treatment these alloys possess many favorable physical and mechanical properties

that make them excellent implant materials (Wataha 1996).

Commercially pure titanium (CP) comes in different grades, from CP grade Ι to CP grade

ΙV, which vary mostly in the oxygen content. The compositions in weight percentage of

these metals and of the two types of titanium alloys that are commercially used in implant

dentistry, as given in several ASTM (American Society for Testing and Materials)

Standards, appear in Table 3. Some of these materials can be supplied in the ELI

condition (Exra Low Interstitial content) (ASTM 2000).

Titanium N C H Fe O Al V Ti

CpTi,Grade 1 0.03 0.10 0.015 0.02 0.18 _ _ balance




Ti 6Al 4V 0.05 0.08 0.015 0.30 0.20 5.50-6.75 3.50-4.50 balance

Ti 6Al 4V(ELI) 0.05 0.08 0.012 0.10 0.13 5.50-6.50 3.50-4.50 balance

Table 3 Titanium grades 1-4 and titanium alloys (Ti 6Al 4V) compositions (weight percent) from

ASTM Standard (ASTM 2000)


Nowadays titanium is applied in all fields of dentistry. Due to its advantages, it is

considered to be a valuable future alternative to conventional dental casting alloys and is

used in various dental applications (Wang and Fenton 1996).

2.5.1 Mechanical properties of cp titanium and alloys

A comparison between the mechanical properties of pure titanium, its heat-treated alloys

and other natural and implant materials is made in Table 4. These properties may or may

not vary with the grade and the alloy type. It is important to note that while the modulus

of elasticity of cp grade 1 titanium to cp grade 4 titanium ranges from 102 to 104 GPa (a

change of only 2%), the yield strength increases from 170 to 483 MPa (a gain of 180%).

This is mainly related to the oxygen residuals in the metal. The characteristic trend of

increasing strength with relatively constant modulus continuous when comparing cp

titanium with its alloys (McCracken 1999).

Compared with Co-Cr-Mo alloys, titanium alloy is almost twice as strong and has half

the elastic modulus. Compared with 316L stainless steel, the Ti 6Al 4V alloy is roughly

equal in strength, but again, it has half the modulus of elasticity (McCracken 1999).

Strength is beneficial because implant materials better resist occlusal forces without

fracture. Lower modulus is desirable because the implant biomaterial better transmits

forces to the bone. Although titanium alloys are stiffer than bone, their modulus of

elasticity is closer to bone than any other important implant metal; the only exception is

pure titanium. That leads to a more even distribution of stress at the critical bone-implant

interface because the bone and implant will flex in a more similar fashion (Parr et al.

1985). Material Elastic Modulus

(GPa)

Tensile Strength

(MPa)

Yield Strength

(MPa)

Elongation

(%)

Cp grade 1 Ti 102 240 170 24

Cp grade 2 Ti 102 345 275 20

Cp grade 3 Ti 102 450 380 18

Cp grade 4 Ti 104 550 483 15

Ti 6Al 4V ELI 113 860 795 10


Ti 6Al 4V 113 930 860 10

Co-Cr- Mo 240 700 450 8

316L steel 200 965 690 20

Cortical bone 18 140 n/a 1

Dentin 18.3 52 n/a 0

Enamel 84 10 n/a 0

Table 4 Mechanical properties of pure titanium and of its alloys (Wataha 1996; ASTM 2000)

Titanium is quite light in weight, having a density of 4.5 g/cm3, which is considerably

less than that of Co-Cr or 316L stainless steel (8.5 and 7.9 g/cm3 respectively). Thus, the

combination of high strength and low weight makes titanium and its alloys some of the

highest strength/weight ratio and mostly preferred current materials in implant dentistry

(Lautenschlager and Monaghan 1993).

2.5.2 Corrosive behavior and biocompatibility of pure titanium and its

alloys

Corrosion leads to the release of compounds into biological environments, which may

then cause adverse effects, such as toxic or allergic reactions. Corrosion resistance is,

therefore, a prerequisite for biocompatibility (Schmalz and Garhammer 2002;

Tschernitschek et al. 2005). Titanium and its alloys generally exhibit biocompatibility and a good corrosion

resistance, as a result of the passivating effect through a thin layer of titanium oxides

(TiO, TiO2, and Ti2O3) that is formed on their surface, when they are exposed to air,

water or any electrolyte (Kappert 1994). Titanium has the ability to form an oxide layer

of 10 Å (1 Nanometer) in thickness within a millisecond and is generally self-healing. If

left uncontrolled, this oxide layer becomes 100 Å thick within a minute (Kasemo 1983).

It consists mainly of TiO2, which has been shown to be stable (Parr et al. 1985;

McCracken 1999). Its role is to separate the reactive titanium from the electrolyte and to


minimize ion release, thus serving as a protective layer, which prevents direct contact of

the metal with the surrounding structures (Wang and Fenton 1996).

The general perception of titanium’s corrosion resistance and biocompatibility have been

questioned somewhat during recent years (Tschernitschek et al. 2005). Titanium was

considered to be the “bioinert metal per se” in the 1980s and 1990s (Lautenschlager and

Monaghan 1993; Kappert 1994). It was thought to be absolutely resistant to corrosion,

and would not cause sensitization or toxic effects. This opinion was supported by the fact

that a sound human body contains approximately 0.01 g Ti/70 kg body weight, which

indicates that titanium is a natural component of human tissues (Wirz and Will 1999).

However, current data have shown that titanium can be corrosive as many other base

metals under mechanical stress, oxygen deficit, or at a low pH level (Tschernitschek et al.

2005). In particular, fluoride reveals a high affinity to titanium. Fluoride ions can

infiltrate and dissolve the stabilizing oxide layer (Hösch and Strietzel 1994). It has been

shown that caries-preventive fluoride gels increase the corrosion and surface roughness

of titanium due to their pronounced adhesion (Lenz 1997). It should be emphasized at

this point that the combination of fluoride concentration and pH value is of great

importance, as at a neutral pH toothpaste with a low concentration of fluoride (0.02% to

0.19%) do not cause relevant corrosion (Wikidal and Geis-Gerstorfer 1999). In the oral

environment, when the pH is almost neutral (pH=7) the passive layer dissolves at such a

slow rate that the resultant loss of mass is of no consequence for the implant. It has been

found that cast titan restorations are more susceptible to corrosion than are machined titan

restorations (i.e. dental implants) (Patyk and Ohm 1997). Koike and Fujii have also

observed that acids in the absence of fluoride ions can cause electrochemical corrosion

(Koike and Fujii 2001).

Electrochemical tests have been used to measure potential corrosion and release of

elements when commercially pure titanium is coupled with other dental alloys, as might

occur when the implants are restored with prosthetic devices. It appears that most dental

alloys pose little threat for causing increased titanium corrosion. However, the risk of

such electrochemical corrosion was significantly higher when titanium was coupled with

non-precious (Ni-based) alloys (Reclaru and Meyer 1994).


Contradictory data have been reported about potential toxic effects and interactions of

titanium and its alloys with various tissues. The normal level of titanium in human tissue

is 50 ppm. Values of 100 to 300 ppm are frequently observed in soft tissues surrounding

titanium implants. At these levels, tissue discoloration with titanium pigments can be

seen. According to some authors this rate of dissolution is one of the lowest of all

passivated implant metals and seems to be well tolerated by the body (Parr et al. 1985).

Cell culture studies revealed no indication of cyto- or genotoxic effects due to the

corrosion products of titanium (Rae 1986a, b; Berstein et al. 1992). On the other hand, a

study by Wirz and Will reported 65 failures of Ti implants, where more than 50% of

those cases were due to eventually toxic metal ions, which were liberated from the

superstructure (Wirz and Will 1999). Another study reported that small titanium particles,

which are generated due to wear, induce programmed cell death, or apoptosis, in

mesenchymal stem cell (Kumazawa et al. 2002). Sensitization to titanium has been

observed with immunologic in-vitro tests [lymphocyte-transformation-test (LTT)] by

other investigators (Lalor et al. 1991; Valentine-Thon and Schiwara 2003). Current

available literature also indicated a risk for titanium (among other metals too)-induced

autoimmunity and non-specific immunomodulation in man, showing that the metal

pathology may be due to toxic or allergic mechanisms. According to the authors the main

factors decisive for disease induced by metals are exposure and genetics which determine

the individual detoxifying capacity and sensitivity to metals (Stejskal and Stejskal 1999).

The transport of titanium away from the implant site has also been documented. Titanium

has been detected at elevated levels in lungs, kidneys, and liver of miniature pigs after the

insertion of screw implants (Schliephake et al. 1989; Schliephake et al. 1993a;

Schliephake et al. 1993b). Weingart et al. investigated the deposition of titanium in

regional lymph nodes after insertion of endosseous, plasma-spray-coated titanium screw

implants in a total of 19 beagle dogs (Weingart et al. 1994). Bianco et al. documented

elevated titanium serum and urine levels in the presence of Ti-based prostheses. However

the authors found no increase in the titanium levels in serum and urine of test animals

with implants (rabbits) in comparison to controls up to 1 year after implantation (Bianco

et al. 1996). Contradictory studies have shown no elevated levels of titanium in the above

mentioned human tissues after two years of implantation (Lugowski et al. 1991).


It is apparent that the dental literature concerning the interactions of titanium and its

alloys with human body is controversial. Although the available information of the above

mentioned studies indicate that titanium can corrode in the oral cavity, as well as in other

tissues, further experiments, specifically clinical studies, are necessary to elucidate these

important aspects, as the adverse effects due to titanium and its alloys are rare and

considerably less pronounced in comparison to that of other metals. However, as the

biostability of titanium is becoming increasingly questioned and at the same time new

technologies and materials, such as high performance ceramics (i.e. zirconium dioxide),

are emerging, the probability arises that these non-metallic materials could eventually

replace metals and alloys in the future.

2.6 The use of zirconium dioxide in dentistry

Due to an increasing interest in esthetics and concerns about toxic and allergic reactions

to certain alloys, patients and dentists have been looking for metal-free tooth-coloured

restorations. Therefore, the development of new high strength dental ceramics, which

appear to be less brittle, less limited in their tensile strength, and less subject to time-

dependent stress failure, has dominated in the later part of 20th century (Qualtrough and

Piddock 1997; Strub and Beschnidt 1998; McLean 2001; Tinschert et al. 2001b).

Zirconium dioxide (zirconia) is a high performance ceramic used so far in dentistry for

fabricating endodontic posts, crown/bridge restorations and implant abutments. It has

been also applied for the fabrication of esthetic orthodontic brackets (Keith et al. 1994).

Zirconia posts with custom-made ceramic cores have been investigated and are already

used for retaining all-ceramic restorations (Meyenberg et al. 1995; Fischer et al. 1998;

Butz et al. 2001; Strub et al. 2001; Heydecke et al. 2002; Jeong et al. 2002; Oblak et al.

2004). Kern et al. in their study followed 80 devital teeth that were restored with zirconia

posts over an average observation period of 16 months and they reported survival rates of

100% (Kern et al. 1998). In another study, Paul and Werder examined 145

endodontically treated teeth restored with zirconia posts. 79 posts with direct composite

cores and 34 posts with glass-ceramic cores were reevaluated clinically and

radiographically after a mean clinical service of 4.8 years and 3.9 years respectively. In a


best-case scenario posts that could not be reevaluated (8 for the first group and 24 for the

second group) were considered successful, and in a worst-case scenario they were

considered failures; respective success rates were 100% and 91% in the direct group and

95% and 53% in the indirect group. However, the authors reported that the high dropout

rate for the second patient group precluded valid conclusions for this type of restoration

(Paul and Werder 2004).

Although still in the experimental stage, the fabrication of zirconia frameworks of either

presintered or highly isostatic pressed zirconia for crown-and-bridge work seems to be

possible (Luthardt and Musil 1997; Luthardt et al. 1999; Tinschert et al. 2001a). Zirconia

frameworks offer new perspectives in metal free fixed partial dentures (FPDs) and single

tooth reconstructions because of zirconia’s high flexural strength of more than 900 MPa

and show good first clinical results. In a clinical study, Sturzenegger et al. examined 22

zirconia bridges fabricated with the DCM-System (Direct Ceramic Machining System).

The reported success rate was 100% for an observation period of one year (Sturzenegger

et al. 2000). A similar success rate (100%) was reported by Rinke, who examined 22

zirconia bridges fabricated with a CAM system (Cercon®, DeguDent, D-Hanau) for a

mean clinical evaluation period of one year (Rinke 2004). In another clinical study, 58

zirconia bridges fabricated by the DCM system exhibited a survival rate of 84% in a

period of 3.5 years. Minor porcelain chipping was reported in 11% of the bridges (Sailer

et al. 2003). Tinschert et al. observed 65 zirconia bridges fabricated with the use of DCS-

President® system (DCS Dental AG, CH-Allschwil) for a mean period of three years.

They reported a small chipping of the veneering material in 6% of the bridges, which

showed a cumulative survival rate of 86% (Tinschert et al. 2005). Other investigators

showed a 15% rate of minor chipping after a two-year observation period of 23 zirconia

FPDs (Vult von Steyern et al. 2005). No fracture of the zirconia substructures was

reported in the above mentioned studies.

In implant dentistry abutments made of densely sintered yttrium-stabilized zirconia have

been introduced to support implant-supported single-tooth crowns and are nowadays in

use. These abutments are distinguished by their tooth-matched colour, their good tissue

compatibility, and their lower plaque accumulation (Studer et al. 1996a, b; Wohlwend et

al. 1996; Wlochowitz et al. 1998; Yildirim et al. 2000; Scarano et al. 2003; Tan and


Dunne 2003; Yildirim et al. 2003). Yildirim et al. compared in their in-vivo study 51

aluminium oxide abutments with 30 zirconia abutments and found cumulative survival

rates of 98.1% and 100% for each group of implant abutments respectively for an

observation period of 28 months (Yildirim et al. 2003). In a prospective study with an

observation period of 4 years, Glauser et al. showed also a cumulative survival rate of

100% for 53 zirconia abutments (Glauser et al. 2004).

Zirconia implants have been used only experimentally so far (Albrektsson et al. 1985;

Akagawa et al. 1993; Akagawa et al. 1998; Dubruille et al. 1999; Sennerby et al. 2005).

Kohal and Klaus presented in the literature the first clinical case of an all-ceramic

custom-made zirconia implant-crown system, which was used for the replacement of a

single tooth (Kohal and Klaus 2004). Volz and Blaschke also published the case report of

a patient with metal sensitivities who received eight custom-made zirconium dioxide

implants restored with metal-free zirconia crowns (Volz and Blaschke 2004). In both

cases a successful osseointegration was obtained.

The current laboratory and clinical trials, regarding zirconia implants, as well as zirconia-

based all-ceramic crowns and FPDs, are encouraging and promising so far, presenting

that this new ceramic material could offer the optimal basis for an esthetical restoration of

missing teeth.

2.6.1 Zirconium and the mineral Zircon

Zirconium is a metallic element that has the symbol Zr and atomic number 40. It

belongs to the transition elements of the periodic table and has an atomic weight of

91.224 (Stevens and Hennike 1992).

The name of the metal zirconium origins from the Arabic “zargun” (golden in colour)

which in turn comes from the two Persian words Zar (gold) and Gun (colour). Zirconia,

the metal dioxide (ZrO2), was identified as such in 1789 by the German chemist Martin

Heinrich Klaproth in the reaction product obtained after heating some gems (Piconi and

Maccauro 1999). The element was isolated first by the Swedish chemist Baron Jöns

Jacob Berzelius in 1824. Pure zirconium was not prepared until 1914.


In its pure state zirconium exists in two forms: the crystalline form, a soft, white, ductile

metal; and the amorphous form, a bluish-black powder. Zirconium ranks 18th in

abundance among the elements in the crust of earth and is never found in nature as a free

metal. It melts at about 1852oC (about 3362oF) and boils at about 4377oC (about 7911oF)

(Stevens and Hennike 1992).

The principal economic source of zirconium is the zirconium silicate mineral, zircon

(ZrSiO4) (Fig. 3), which is found in deposits located in Australia, Brazil, India, Russia,

and the United States. Australia is the largest producer of zirconium in the world,

accounting for more than 70 percent of world production. Zirconium occurs also as an

oxide (ZrO2) in the mineral baddeleyite (Fig. 4) and is recognized in 30 other mineral

species. Commercial – quality zirconium still has a content of 1 to 3% hafnium, a metal

with properties similar to those of zirconium (Piconi and Maccauro 1999).

Fig. 3 The mineral Zircon (ZrSiO4) Fig. 4 The mineral Baddeleyite (ZrO2)

Zircon is a transparent, translucent, or opaque mineral that may occur as colorless

crystals or in shades of green, gray, red, blue, yellow or brown. The clear, transparent

varieties are often used as gemstones and are known as hyacinth or jacinth; translucent or

opaque varieties are known as jargon. Zircon is often found together with gold, as

rounded grains in streams and along sandy beaches

Zircon contains zirconium and hafnium at a ratio of about 50 to 1. They are difficult to

separate.


The major end uses of zircon are refractories, foundry sands, and ceramic opacification.

Zircon is also marketed as a natural gemstone used in jewelry, and its oxide is processed

to produce the diamond stimulant, cubic zirconia (Stevens and Hennike 1992).

2.6.2 The mineral zirconium dioxide

Zirconia or zirconium dioxide (ZrO2) is the white crystalline oxide of zirconium. The

metal zirconium is commercially produced from Zirconia in the Kroll Process in a similar

way to titanium. The production involves the action of chlorine and carbon upon

baddeleyite. The resultant zirconium tetrachloride, ZrCl4, is separated from the iron

trichloride, FeCl3, by fractional distillation. Finally, ZrCl4 is reduced to metallic

zirconium by reduction with magnesium. Air is excluded so as to prevent contamination

of the product with oxygen or nitrogen.

Although low-quantity zirconia is used as an abrasive in huge quantities, tough, wear

resistant, refractory zirconia ceramics are used to manufacture parts operating in

aggressive environments, like extrusion dyes, valves and port liners for combustion

engines, low corrosion, thermal shock resistant refractory liners of valve parts in

foundries (Stevens and Hennike 1992).

Most often zirconium does not occur in nature in a pure state. Its two forms (baddeleyite

or zircon) are commonly associated with rutile, ilmenite and monazite with significant

concentrations of natural radionuclides, like uranium and thorium.

For the use of zirconium dioxide as a high-performance ceramic material, the zirconia

powder production processes (from baddeleyite or zircon) perform an effective separation

of elements, such as titanium and iron from baddeleyite, and SiO2 from zircon.

Nevertheless, uranium, thorium and their decay products can be present at impurity levels

in some zirconia powders. Their concentrations depend on the powder production process

and on the purification level attained. The presence of such impurities can be disregarded

in ceramics to be used as refractories or as combustion engine parts, but it has to be

carefully assessed in ceramic biomaterials (Piconi and Maccauro 1999; Strub et al. 2005).


2.6.3 Zirconium dioxide as a high-performance ceramic biomaterial

The origin of the interest in using zirconia as a ceramic biomaterial was its good chemical

and dimensional stability, its mechanical strength and toughness, coupled with a Young’s

modulus in the same order of magnitude of stainless steel alloys. It was introduced into

medicine and dentistry as an ideal replacement for metal (Piconi and Maccauro 1999).

The research on zirconia as a biomaterial was started in the late sixties. Zirconium

dioxide is used in medicine for the manufacturing of artificial hips-, finger- and acoustic-

implants. Christel et al. introduced the first paper concerning the use of zirconia to

manufacture ball heads for Total Hip Replacements, which is the current main

application of this ceramic biomaterial in medicine (Christel et al. 1989).

In the early stages of the development several solid solutions (ZrO2-MgO, ZrO2-CaO,

ZrO2-Y2O3) were tested for biomedical applications. But in the following years the

research efforts appeared to be more focused on zirconia-yttria ceramics, characterized by

fine-grained microstructures known as Tetragonal Zirconia Polycrystals (TZP) (Cales et

al. 1994; Drouin et al. 1997; Piconi et al. 1998; Piconi and Maccauro 1999).

Zirconia is a well-known polymorph that can exist in three metamorphs (phases) termed

monoclinic (m), tetragonal (t) and cubic (c). Pure zirconia has a monoclinic crystal

structure at room temperature. This phase is stable up to 1170oC. Above this temperature

there is a transition into the tetragonal and then into the cubic phase at 2370oC. During

cooling a t→m transformation takes place at a temperature range of about 100oC below

1070oC. However, noticeable changes in volume are associated with these

transformations: during the monoclinic to tetragonal transformation a 5% volume

decrease occurs when zirconium oxide is heated; reversely, a 3-4% increase in volume is

observed during the cooling process (Fig. 5) (Christel et al. 1989). Stresses generated by

the expansion lead to cracks in pure zirconia ceramics that can break into pieces at room

temperature. The transformation into the monoclinic phase is of the same nature as the

martensitic transformation occurring in steel and can be compared to it (Garvie et al.

1975; Piconi and Maccauro 1999).


Up to 1170oC 1170oC-2370oC 2370oC-2680oC 5% Volume decrease 3-4% Volume increase

Fig. 5 Temperature-related phase transformation of zirconium dioxide

Several different oxides are added to zirconia to stabilize the tetragonal and/or cubic

phases. Magnesia (MgO), Yttria (Y2O3), Calcia (CaO), and Ceria (CeO), amongst others,

allow to generate multiphase materials known as Partially Stabilized Zirconia (PSZ),

whose microstructure at room temperature generally consists of cubic zirconia as the

major phase, with monoclinic and tetragonal zirconia precipitates as the minor phase

(Christel et al. 1989; De Aza et al. 2002). In 1972, Garvie and Nicholson showed that the

mechanical strength of PSZ was improved by an homogenous and fine distribution of the

monoclinic phase within the cubic matrix (Garvie and Nicholson 1972). In the presence

of a small amount of stabilizing additive, tetragonal particles, provided they are small

enough, can be maintained in a metatastable state, at temperatures below the t→m

transformation temperature. The transformation of small tetragonal grains, which should

result in a volume increase, is prevented by the compressive stresses applied on these

grains by their neighbours. In the ZrO2-MgO or ZrO2-CaO systems, materials are sintered

in the cubic state and small tetragonal precipitates are formed during the cooling as a

result of partial transformation of the cubic phase.

In the ZrO2-Y2O3 system, it is also possible to obtain ceramics formed at room

temperature with only a tetragonal phase, called Tetragonal Zirconium Polycrystal (TZP)

(Christel et al. 1989). Thus, using Y2O3 as a stabilizing agent, it is possible to produce a

zirconium dioxide ceramic made of 100% small metatastable tetragonal grains. A

propagating crack can release the stresses on the neighbouring grains, which then

transform from the metastable tetragonal state into the monoclinic phase (Fig. 6). 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. This phase transformation

Monoclinic Phase

Tetragonal Phase

Cubic Phase


can put the crack into compression (because of the size of the grains), retarding its

growth. An enhancement in toughness is obtained, because the energy associated with

crack propagation is dissipated both in the t→m transformation and in overcoming the

compression stresses due to volume expansion. This mechanism is known as

transformation toughening, and is considered to be the basis for the very high strength

of yttria-tetragonal zirconia polycrystal (Y-TZP) (Garvie et al. 1975; Gupta et al. 1978;

Thompson and Rawlings 1990; Cales 2000; Clarke et al. 2003; Guazzato et al. 2004).

Fig. 6 An advancing crack front is dissipated in zirconia as a result of the energy

required to (a) transform the tetragonal phase to the monoclinic phase and (b) overcome

the compressive stress field of the expanded monoclinic grain (Clarke et al. 2003).

2.6.4 Mechanical properties and ageing of Y-TZP There is no doubt that zirconia ceramics have mechanical properties better than all other

ceramic biomaterials. In Table 5 some of the most important mechanical properties of

Surgical Grade Alumina and Y-TZP either pressureless sintered or sintered and high

isostatically pressed are presented (Christel et al. 1989). A high isostatic pressing (HIP)

cycle after sintering is recommended to the manufacturing process of zirconium dioxide

to reach full density that is close to the theoretical (d= 6.1=100% dth).


Property Alumina Y-TZP (pressureless sintering)

Y-TZP (Sinter + HIP)

Density (g/cm3) 3.90 6 6.1

Average grain size

(μm)

<7 <1 <0.5

Microhardness

(Vickers)

2000-3000 1000-1200 1000-1300

Young’s Modulus

(GPa)

380 200 200

Bending Strength

(MPa)

400 900 1200

Toughness KIC

(MPa·m1/2)

5-6 9-10 9-10

Table 5 Comparison of several mechanical properties between alumina and yttria stabilized

zirconia (Y-TZP) either pressureless sintered or sintered and high isostatically pressed

The mechanical properties of zirconia relate to its fine-grained, metastable

microstructure. The stability of this structure during the lifetime of TZP components is

the key point to attain the expected performances.

Under certain manufacturing conditions or more severe environmental conditions of

moisture and stress, the resulting zirconia may transform more aggressively to the

monoclinic phase with catastrophic results. Such a “high metastability” is obviously

undesirable for medical implants. This mechanical property degradation in zirconia, due to the progressive spontaneous transformation of the metastable tetragonal phase into the

monoclinic phase, is known as “ageing” of the material (Cales et al. 1994; Cales 2000;

Ardlin 2002).

A low-temperature degradation, which has a maximum rate at 250o C, has been studied in

detail (Matsumoto 1985; Sato and Shimada 1985). According to Swab (1991), the

transformation is enhanced in water or in vapour, while the most critical enhancing

effects of temperature are in the range 200-300oC. The t→m transition starts from the

surface and progresses into the material bulk. Resistance to transformation is increased


by a small grain size (<<1 μm), a density as close to the theoretical (d=6.1 g/ cm3) and

yttrium oxide content as close as possible to 3 moles percent (5.1 weight percent).

The increase in monoclinic phase leads to a reduction in strength, toughness and density,

followed by micro-macrocracking (Swab 1991). The growth of the transformed zone

leads to extensive micro cracking and surface roughening. In aqueous environments, this

offers a path for the water to penetrate down into the specimen, creating corrosion effects

on the Zr-O-Zr bonds (Sato and Shimada 1985; Chevalier 2006). The growth stage

depends on several microstructural patterns: porosity, residual stresses, grain size, etc

(Deville et al. 2006). Reduction in grain size and/or increase in concentration of

stabilising oxides can reduce the transformation rate. However reducing the size of grains

too much, may lead to the loss of metastability, and increasing the concentration of

stabilizing oxide above 3.5 mol% may allow the nucleation of significant amounts of the

stable cubic phase (Gupta et al. 1978; Theunissen et al. 1992).

The Y-TZP implant ceramic was studied extensively in this regard and had been

considered to be stable under normal body conditions by several authors (Christel et al.

1989; Ichikawa et al. 1992; Shimizu et al. 1993; Cales et al. 1994; Piconi et al. 1998;

Piconi and Maccauro 1999), until the year 2001 when roughly 400 femoral heads of the

Prozyr® product failed in a very short period, as a result of accelerated ageing of the

material (Chevalier 2006). The results depended on the quality and microstructural design

of the zirconia used, as all zirconias are not the same. Even if limited in time and number,

and clearly identified to be process controlled, these events have had a bad impact for the

use of zirconia in the orthopaedic community, resulting in a reduce of market sale of

more than 90% between 2001 and 2002 regarding zirconia femoral heads (Chevalier

2006). The controversial results about the performance of zirconia as a biomedical

material highlight the need of zirconia biomaterials to be manufactured according to the

guidelines of the existing ASTM (American Society for Testing and Materials) and ISO

(International Organization for Standardization) standards.


2.6.5 Biocompatibility of Y-TZP

Zirconia ceramics are chemically inert materials and demonstrate a high biocompatibility.

Piconi and Maccauro have reported a detailed review of biocompatibility of zirconia

ceramics for in vitro and in vivo tests (Piconi and Maccauro 1999).

The in vitro tests generally concluded that zirconia ceramics have no cytotoxic effect on

fibroblasts. Covacci et al. demonstrated in their study that purified Y-TZP has no oncogenic and mutagenic effects in vitro and it can be considered suitable for biomedical

applications (Covacci et al. 1999).

In vivo short-term and long-term effects also were investigated and there is a general

agreement on the absence of local or systemic effects after implantation of zirconia into

muscles or bones (Piconi and Maccauro 1999). Christel et al. reported on the in vivo

behaviour of a Y-TZP ceramic after short (3 months) or mid-term (6 months)

implantation time in rat paraspinal muscle and rabbit leg bones. In both cases the in vivo

effect of the zirconia was compared with that of alumina and no difference was detected

between the two ceramics as the biocompatibility concerns (Christel et al. 1989).

Ichikawa et al. also evaluated the tissue reaction and stability of partially stabilized

zirconia ceramic in vivo with the use of the subcutaneous implantation test in rats.

Histological findings in this study (3, 6 and 12 months after the implantation) indicate

that zirconia ceramic has been well tolerated in the subcutaneous tissue, encapsulated by

a thin fibrous connective tissue, comparable to polycrystalline alumina, which is reported

to be tissue compatible. The results suggested that zirconia ceramics are highly

biocompatible (Ichikawa et al. 1992).

Comparison studies between titanium and zirconia ceramic implants that were inserted in

the bone of animals give also very encouraging results, concluding that both kinds of

implants osseointegrate to the same extent, being well accepted, as indicated by the lack

of adverse tissue reactions to the implants and the high level of direct contact between

bone and both materials, titanium and zirconia (Albrektsson et al. 1985; Scarano et al.

2003; Kohal et al. 2004). Akagawa and colleagues compared the bone tissue response to

loaded and unloaded zirconia implants in the monkey. The authors showed that partially

stabilized zirconia implants placed with a one-stage procedure achieve long-term stability


of osseointegration. He used in his experiment single freestanding implants, connected

freestanding implants, and implant-tooth supports (all implants were immobile after 24-

month of loading) (Akagawa et al. 1998). No mechanical problems, such as fracture of

the implants were reported, indicating the favourable mechanical properties of zirconia,

which is in line with the experiences of Kohal and colleagues. These authors compared

custom-made titanium and zirconia implants used to support metal crowns in the maxilla

of six monkeys. The implants were allowed to heal for 6 months prior to abutment

connection. Metal crowns were cemented after 3 months, and the animals were followed

for another 5 months. All implants achieved and maintained stability, and no mechanical

problems were reported (Kohal et al. 2004). Similarly, Sennerby et al. in a study in the

rabbit showed a strong bone tissue response to surface-modified zirconia implants after 6

weeks of healing, confirming the biocompatibility of this new material (Sennerby et al.

2005).

However little information is available on the resistance of dental implants towards load.

Only one Finite Element Analysis paper can be found in the international literature

(Kohal et al. 2002). Therefore, the present investigation aimed to evaluate the stability of

a one-piece zirconia implant system.

Aim of the study 30

3. Aim of the study

The aim of this in-vitro study was to evaluate the survival rate, after exposure to the

artificial mouth, and the fracture strength of unrestored and restored one-piece ZrO2

implants. The results were compared to those of titanium implants. The ceramic implants

were fabricated out of yttria-stabilized tetragonal zirconia polycrystal (Y-TZP) with

different surface topographies and abutment preparation designs.

Materials and Methods 31

4. Materials and methods

4.1 Outline of the study (Fig.7)

A total of one hundred twenty screw type ceramic and titanium implants were used for

the experiment. The two control groups A and B included sixteen titanium implants each.

The ceramic implants were divided into five groups C, D, E, F and G using sixteen

samples of each, with the exception of group D, in which twenty-four samples were used.

The test and the control groups were divided into 15 subgroups of 8 samples each. Each

group included two subgroups, with the exception of group D that included 3 subgroups.

Groups A and B consisted of two different kinds of titanium implants: group A

comprised a two-piece implant system and group B a one-piece implant system. Groups

C and D included zirconium dioxide implants fabricated from two different companies.

The implants of groups E, F and G had a special surface topography (ZiUnite). The

implant heads of group F were prepared with a 0.5 mm chamfer. The implant heads of

group G were prepared in two different ways: subgroup G1 with a 0.5 mm chamfer and

G2 with a 1 mm chamfer respectively. Subgroups G1 and G2 were restored with all-

ceramic crowns. Subgroups A2, B2, C2, D2, E2, F2, G1, G2 were exposed to 1.2 million

loading cycles including thermocycling in the artificial mouth. Subgroup D3 was exposed

to the artificial mouth without thermocycling. Then, all test specimens (exposed and

unexposed to the artificial mouth) were loaded in a universal testing machine until

fracture occurred.


120 Dental Implants

Control Group A Brånemark ® Ti implants (16)

Control Group B Nobel Direct ® Ti implants (16)

Test Group C Y-TZP BIO- HIP Sigma® implants (16)

Test Group D Uncoated Y-TZP-A BIO- HIP® implants (24)

Test Group E Y-TZP-A BIO-HIP Ziunite ® impl. (16)

Test Group F Y-TZP-A BIO-HIP Ziunite ® implants 0.5 mm chamfer (16)

Test Group G Y-TZP-A BIO-HIP Ziunite ® implants with crowns (16)

A1 (8)

A2 (8)

B1 (8)

B2 (8)

C1 (8)

C2 (8)

D1 (8)

D2 (8)

D3 (8)

E 1 (8)

E2 (8)

F1 (8)

F2 (8)

G1 (8) 0.5 mm

G2 (8) 1 mm

Exposure to the artificial mouth (thermocycling

Exposure to the artificial mouth (thermocy-cling)


Exposure to the artificial mouth (thermocycling)

Exposure to the artificial mouth (no thermocycling)


Exposure to the artificial mouth (thermocycling)

Fracture resistance test

Fig. 7 Outline of the study


4.2 Materials used for the fabrication of the implants

The following materials have been examined:

Group A

• 16 Brånemark titanium implants (3.75 mm x 15 mm, external hex) restored with a

custom-made abutment; attached with a torque of 32 Ncm (Nobel Biocare AB, SE-

Gothenburg).

Control groups A1 (no artificial load, fracture testing) and A2 (with artificial load, fracture

testing).

Group B

• 16 one-piece NobelDirect titanium implants (3.75 mm x 15 mm) (Nobel Biocare AB,

SE-Gothenburg).

Control groups B1 (no artificial load, fracture testing) and B2 (with artificial load, fracture

testing).

Group C

• 16 yttria reinforced - hot isostatically pressed tetragonal zirconia polycrystal (Y-TZP

BIO-HIP) Sigma implants (4.28 mm x 14.4 mm) (Incermed, CH-Lausanne).

Test groups C1 (no artificial load, fracture testing) and C2 (with artificial load, fracture

testing).

Group D

• 24 yttria reinforced - hot isostatically pressed tetragonal zirconia polycrystal A (Y-

TZP-A BIO-HIP) implants (4.3 mm x 16 mm) (Metoxit AG, CH-Thayngen).

Test groups D1 (no artificial load, fracture testing), D2 (with artificial load, fracture

testing), and D3 (with artificial load, without thermocycling, fracture testing).

Group E


TZP-A BIO-HIP) implants (4.3 mm x 16 mm) with ZiUnite surface (Nobel Biocare AB,

SE-Gothenburg).


Test groups E1 (no artificial load, fracture testing) and E2 (with artificial load, fracture

testing).

Group F



SE-Gothenburg).

Test groups F1 (0.5 mm chamfer preparation, no artificial load, fracture testing) and F2 (0.5

mm chamfer preparation, with artificial load, fracture testing).

Group G



SE-Gothenburg).

Test groups G1 (0.5 mm chamfer preparation, all-ceramic crown, with artificial load,

fracture testing) and G2 (1 mm chamfer preparation, all-ceramic crown, with artificial load,

fracture testing).

4.2.1 Y–TZP BIO-HIP Sigma® Implants (Incermed, CH-Lausanne)

• Manufacturing Process

This material is a ceramic-based zirconium reinforced by thermal and physical

processing. Its resistance has been especially increased by a technique of regular ranking

of the crystals binding them with proper materials.

The crystalline products are processed into powder by grinding and compressed by an

isostatic process at high temperature up to about 2000oC.

The pure powder of zirconium, for which the granulometric spectrum has been defined, is

processed through pressure in high temperature molds, resulting in homogenous implants

of exact dimensions.


• Properties

The material properties of HIPped (high isostatically pressed) Y- TZP are summarised in

Table 6.

Property Unit Y-TZP BIO-HIP®

Chemical Composition

ZrO2+HfO2+ Y2O3

Y2O3

Other Oxides

% mass fraction

> 99.9

5.2

< 0.1

Bulk Density g/cm3 6.09

Microstructure

Mean linear intercept distance

μm

0.2

Microhardness (Vickers) HV 1200

Biaxial flexure strength MPa 900

Compressive strength MPa 1650

Young’s Modulus GPa 200

Hydrothermal Stability Good

Table 6 Properties of Y-TZP BIO-HIP® (Information obtained from manufacturer)

4.2.2 Y-TZP-A BIO-HIP® Implants (Metoxit AG, CH-Thayngen)

• Manufacturing Process

The main steps of the manufacturing process for dental HIPped Y-TZP-A implants are

outlined in Figure 8. It can be divided into two separate chapters:

1. Material compaction, sintering, HIPping, reoxidizing

2. Machining (grinding), measurement proof testing and quality control.


Especially the first chapter defines the material properties and the mechanical

performance of the implant. For the safety and reliability of the resulting dental implant it

is very important that these steps are carried out in a strictly controlled and validated

procedure:

Raw material formulation

Pressing

Sintering

Grinding

Fig. 8 Manufacturing process of HIPped Y-TZP-A

• Properties

The material properties of HIPped Y-TZP-A are summarised in Table 7. Positive for

dental applications are the strength of 1300 MPa (4 point bending strength) or 850 MPa

(biaxial flexure strength), fracture toughness of 6-10 MPa·m1/2, Weibull Modulus of 14.

The biocompatibility of Y-TZP-A BIO-HIP® has been established in tests required by

ISO 30993-1. As reference material, high purity alumina was used. All tests showed good

results, displaying no reactions or anomalies.

HIP-post-compaction

Oxidizing

Proof testing and

quality control


Property Unit Y-TZP-A BIO-HIP®


ZrO2 (+HfO2)

Y2O3

Al2O3

% mass fraction

>95.0

4.0

0.25

Bulk Density g/cm3 6.06

Microstructure


μm

0.35


Biaxial flexure strength

4-point bending strength

MPa

MPa

850

1300

Compressive strength MPa 2000


Fracture toughness K1c MPa·m1/2 6-10

Weibull Modulus m 14


Colour White

Table 7 Properties of Y-TZP-A BIO-HIP® (Information obtained from manufacturer)

The implants of groups E, F and G (Ziunite® implants, Nobel Biocare, SE-Gothenburg)

were coated with a slurry containing zirconia powder and a pore-former (patent

application SE0302539-2), so that a porous surface could be achieved (Adilstam F,

Iverhed M. Patent SE0302539-2. September 24, 2003). After the coating was applied, the

implants were sintered to full density, under which the pore-former burned off and left a

porous surface (Fig. 9).


Fig. 9 Scanning electron microscopic (SEM) image of the surface structure of a Ziunite® implant

4.3 Implant head preparation

The heads of the ceramic implants were prepared in groups F and G with diamonds used

for the preparation of natural teeth (diamond burs No. 878 012, No.368 023, No. 8878

012, No. 8368 023, Gebr. Brasseler, D-Lemgo). The preparation was carried out as

follows:

Preparation for group F (Fig. 10a and 10b) The implant heads were prepared with an occlusal angle of convergence of about 6o. The

head height was 6 mm, with a shortening of the incisal edge of 2 mm. The cervical finish

was a 0.5 mm deep chamfer preparation.

Preparation for group G In subgroup G1 the implant heads were prepared in the same manner as in group F. In

subgroup G2 the only difference in the preparation form was that the cervical finish was a

1 mm deep chamfer.


Fig. 10a Zirconia implants as machined Fig. 10b Chamfer preparation of a Ziunite®

implant

4.4 Final impressions and fabrication of the crowns

After the preparation of the implant heads in group G impressions were taken with a

polyether impression material (Impregum Penta®, 3M-Espe, D-Seefeld). Then, the

implants were restored with all-ceramic crowns. The crowns were fabricated out of

Procera Zirconia® (Y-TZP) and veneered with Nobel Rondo® (Nobel Biocare AB, SE-

Gothenburg). The Procera® all-ceramic system (Nobel Biocare AB, SE-Gothenburg) is indicated for the

fabrication of all-ceramic (aluminum oxide or zirconium dioxide) crowns, bridges and

abutments (Hegenbarth 1996). The system, which was developed in 1993 (Andersson and

Oden 1993), consists of a computer-controlled scanning and design station located in the

dental laboratory that is connected via a modem to Procera Scandvik AB in Stockholm,

Sweden. The process for the fabrication of the crowns included a scanning of the design

of the abutment with the help of Procera® scanner Mod 50 (Nobel Biocare AB, SE-

Gothenburg). The gathered information was sent to the central Procera® laboratory in

Stockholm, Sweden in order to fabricate the framework. Afterwards, the framework was

veneered, in order to obtain the final restoration. The material properties of Procera

Zirconia® (Y-TZP), as obtained from the manufacturer are summarised in Table 8.


Property Unit Y-TZP


ZrO2+HfO2+ Y2O3

Y2O3

Al2O3

% mass fraction

>99

4.5-5.4

<0.5

Bulk Density g/cm3 >6.05

Microstructure


μm

<0.5


Biaxial flexure strength MPa 1121


Fracture toughness K1c MPa·m1/2 10


Table 8 Properties of Procera Zirconia® (Y-TZP) (Information obtained from manufacturer)

4.5 Luting procedures (Fig. 11) Each implant was cleaned with 96% alcohol. The pre-treatment of the intaglio surface of

the porcelain crowns before cementation involved airborne-particle abrasion with 110 μm

aluminum oxide and a pressure of 2.5 bar for 15 seconds. Excess powder particles were

removed by a gentle stream of room air. Subsequently, the intaglio surface was activated

with Clearfil SE Primer® and Clearfil porcelain bond Activator® (Kuraray, J-Osaka)

mixed 1:1.

For the cementation of the all-ceramic crowns the self-curing composite Panavia® 21

(Kuraray, J-Osaka) was used. Equal amounts of Panavia® 21 Catalyst and Universal

pastes were dispensed and mixed for 20-30 seconds until a smooth, uniform paste

resulted. The paste was applied to the inner surface of the crowns and then the

restorations were carefully seated on the heads and held in place with finger pressure.

Any excess cement was removed and a glycerin gel (Oxyguard II®, Kuraray, J-Osaka)


was applied at the marginal area of the restorations. Afterwards, the gel was removed

with water spray.

Fig. 11 Cementation of a porcelain crown with Panavia 21®

4.6 Fabrication of the master models for the artificial mouth

In order to simulate the clinical situation, all test specimens were embedded using a self-

curing resin (Technovit 4000®, Kulzer, D-Wehrheim). A sample holder of the artificial

mouth was filled with a polyvinyl-siloxane impression material (Putty Soft®, 3M-Espe,

D-Seefeld) (Fig. 12). One implant of each group was embedded into the holder, 1.5 mm

below the finish line of the neck of each implant, in order to simulate the normal bone

resorption that takes place after the first year of the implantation (Adell et al. 1981). The

buccal-lingual inclination angle between the long axis of the implant and the horizontal

plane of the sample holder was 130o replicating the clinical position of anterior teeth (Fig.

13). A silicon mold of the representative model was fabricated seven times (for each

subgroup) with the same impression material (Putty Soft®, 3M-Espe, D-Seefeld) (Fig.14).

Each silicon mold, incorporating this angulation, was used as the negative form for fixing

the implant in the sample holder. The implants were fixed into the silicon mold with the

use of wax (Fig. 15). The sample holder was then isolated with Vaseline (Weißes

Vaselin®, Lichtenstein, D-Mülheim-Kärlich) and attached to the prepared silicone mold

with the implant in place (Fig. 16). Then a self-curing polyester resin (Technovit 4000®,

Kulzer, D-Wehrheim) was mixed and poured into the sample holder from a hole placed


in the lower surface of the sample holder. After the resin had set, the silicone was

removed, and the heads of the implants were cleaned with steam (Fig. 17).

Fig. 12 Sample holder filled with Putty Soft® impression material (3M-Espe, D-Seefeld)

Fig. 13 A Ziunite® implant embedded in Fig. 14 Fabrication of a silicon mold the sample holder (inclination angle 130o)

Fig. 15 The implant fixed into the Fig. 16 The sample holder attached to the

silicon mold with the use of wax silicon mold with the implant in place


Fig. 17 A Ziunite® implant embedded in the sample holder (inclination angle of 130o)

4.7 Experiments 4.7.1 Exposure of the test samples to the artificial mouth

Seventy-two of the specimens (subgroups A2, B2, C2, D2, D3, E2, F2, G1, G2) were

artificially aged in a computer-controlled dual axis-chewing simulator (Willytech, D-

Munich) in order to simulate five years of clinical service (Fig. 18).

The artificial oral environment consists of eight identical sample chambers, two stepper

motors controlling vertical and horizontal movement of the samples against the

antagonist, a hot and cold water circulation system (Haake, D-Karlsruhe) and a

computerized control unit (Fig. 19). The masticatory load curve was programmed by the

combination of the horizontal (0.5 mm) and vertical (6 mm) motion. All of the samples in

each group were subjected to 1.2 million cycles by a reproducible dynamic occlusal load

(Table 9).


Fig. 18 The dual-axis chewing simulator (Willytech, D-Munich)

Fig. 19 Schematic drawing of the dual-axis chewing simulator with eight sample

chambers. (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

removal of cold water, (7) pump for removal of warm water, (8) pump for application of

cold water, (9) pump for application of warm water, (10) motor block, (11) table (Kern et

al. 1999).


Settings of the artificial mouth (Table 9)

Chewing cycle 1,200,000

Cycle frequency 1.6 Hz

Vertical movement 6 mm

Horizontal movement 0.5 mm

Descending speed 60 mm/sec

Rising speed 55 mm/sec

Forward speed 60 mm/sec

Backward speed 55 mm/sec

Applied weight per sample 10 kg (98 N)

Hot dwell time 60 s

Hot bath temperature 55oC

Cold dwell time 60 s

Cold bath temperature 5oC

Intermediate pause 12 s

Table 9 Overview of the parameters in the artificial mouth

Instead of following the standard protocol of applying 49 N load (Krejci and Lutz 1990;

Krejci et al. 1990; Kern et al. 1999), the applied load was 98 N, and the thermo-cycling

was 5oC to 55oC for 60 seconds each with an intermediate pause of 12 seconds,

maintained by the thermostatically-controlled liquid circulator (Haake, D-Karlsruhe).

Each of the eight sample chambers had a plastic sample holder, which was adjusted and

fixed with a butterfly nut to the base of the sample chamber and the underlying lower

crossbeam (Fig. 20). The lower crossbeam was moved by one stepper motor and allowed

a horizontal, sliding motion of the samples (Kern et al. 1999).

Vertical guide rails were freely mounted within the bearings of the upper crossbeam. The

vertical height of the ceramic antagonist Steatit® balls (6 mm in diameter) (Höchst Ceram

Tec, D-Wunsiedel) was adjusted by the adjustment screw on top of the upper crossbeam

(Fig. 18). Weights of 10 kg were mounted on top of the guide rails and established a


chewing force of 98 N. The upper crossbeam was moved by the second stepper motor

and moved the antagonistic steatite balls vertically. Because the guide rails were freely

mounted within bearings in the crossbeam, their individual weight (10 kg) was fully

transferred to each lower sample. The effective impact force is dependent on the

antagonist’s total weight and on its velocity, both of which could be precisely controlled.

The chewing machine’s computer unit calculates and displays the effective impact as

kinetic energy (4,500 x 10-6) (Kern et al. 1999).

Each sample chamber was equipped with a water nozzle through which the sample was

sprayed alternately with cold and warm water (Fig. 20). Two water baths with electronic

temperature and level control supplied cold and warm water of the desired temperature.

To prevent the mixing of cold and warm water, the preceding water was fully suctioned

out before water of the other temperature was applied (Kern et al. 1999).

During the dynamic loading, all samples were examined twice a day. Fractures of the

implants or the crowns (for those groups that were restored) were recorded as a failure.


Fig. 20 Schematic drawing of one chewing chamber. The sample rests on the sample

holder that is fixed to the chamber base by a butterfly nut (Kern et al. 1999).

4.7.2 Fracture strength test (Fig. 21)

All samples that survived the exposure to the artificial mouth were statically loaded using

a universal-testing machine (Zwick, Z010/TN2S, D-Ulm), until fracture occurred.

Each sample was mounted in the test machine. A 0.8-mm thick tin foil (DT Dental

Tradint, D-Bad Kissingen) was placed over the occlusal surface of the implants. The test

started via a computer connected to the testing machine. A perpendicular load was

applied on the angulated implants under a crosshead speed of 2 mm/min at an angle of

130 degrees to the horizontal plane. The loads required for fracturing the samples were

recorded on an X-Y writer with the Zwick testXpert® V 7.1 software, with failure

recorded at the first sharp drop-down of the graphical curve (fracture of the ceramic,


bending of the titanium). The recorded data were automatically analyzed and a graph was

drawn for each sample.

Fig. 21 Fracture strength test of a sample

4.7.3 Statistical analysis of data

The statistical analysis of the data was performed at the Institute of Medical Biometry

and Medical Informatics, Albert-Ludwig University, Freiburg, Germany, using multiple

pair-wise comparisons with the Wilcoxon rank sum test (Splus statistic program, version

3.4 release 1 for Sun SPARC) at a significance level of 0.05. Comparisons for the

following groups were carried out: A1-E1, A2-E2, C1-E1, C2-E2, D2-D3, D1-E1, D2-

E2, E1-E2, E1-F1, E2-F2, F1-F2 and G1-G2.

Results 49

5. Results

5.1 Dynamic loading in the artificial oral environment

5.1.1 Survival rate of the implants

During the dynamic loading of the samples in the artificial mouth, the following results

were observed. There were abrasion traces at the surface of some of the implant

abutments/heads or the porcelain crowns that were in contact with the steatite ball (Fig.

22). This was not considered as failure of the sample.

Fig. 22 Abrasion trace at the surface of an implant crown

As far as the fractures are concerned, seven out of the hundred twenty samples failed in

the chewing simulator.

The following incidences occurred in the different subgroups:

• Control Subgroup A2 (Brånemark® Ti-Implants)

In control subgroup A2, the abutment screw of sample No. 6 fractured at 475,000

chewing cycles (Fig. 23).

Results 50

Fig. 23 Fracture of the abutment screw of a Brånemark®implant

• Test Subgroup C2 (Y-TZP BIO-HIP Sigma® Implants)

In test subgroup C2, four out of the eight samples fractured. A partial fracture of the

implant head of sample No. 5 occurred at 2,000 chewing cycles. The implant heads of

samples No. 1, 2 and 7 fractured at 250,000 chewing cycles (Fig. 24).

Fig. 24 Fracture of the abutment of sample No. 1

• Test Subgroup D3 (Y-TZP-A BIO-HIP ® Implants) In test subgroup D3 (no thermocycling) there was a chipping of the implant head of

sample No. 2 at 190,000 chewing cycles.

• Test Subgroup E2 (Y-TZP-A BIO-HIP Ziunite® Implants)

The head of sample No. 5 of the test subgroup E2 (no abutment preparation of the

implants) fractured at 275,000 chewing cycles.

All other specimens in the subgroups survived 1,200,000 chewing cycles in the artificial

mouth with a load of 98 N.

Results 51

87.5% 100%

50%

100% 87.5% 87.5%

100% 100% 100%

0

20

40

60

80

100

A2 B2 C2 D2 D3 E2 F2 G1 G2

Control Subgroup A2: Brånemark® Ti implants

Control Subgroup B2: Nobel Direct® Ti implants

Test Subgroup C2: Y-TZP BIO-HIP Sigma® implants

Test Subgroup D2: Uncoated Y-TZP-A BIO-HIP® implants (with thermocycling)

Test Subgroup D3: Uncoated Y-TZP-A BIO-HIP® implants (without thermocycling)

Test Subgroup E2: Y-TZP-A BIO-HIP Ziunite® implants (no preparation of the implant heads)

Test Subgroup F2: Y-TZP-A BIO-HIP Ziunite® implants (0.5 mm chamfer)

Test Subgroup G1: Y-TZP-A BIO-HIP Ziunite® implants with crowns (0.5 mm chamfer)

Test Subgroup G2: Y-TZP-A BIO-HIP Ziunite® implants with crowns (1.0 mm chamfer)

Fig. 25 The survival rate of the different subgroups after exposure to artificial mouth

5.2 Fracture strength test

5.2.1 Fracture strength of the implants

The results of the fracture tests in the universal testing machine for each individual

sample can be depicted from the following tables. From a statistical standpoint the

specimens that failed in the artificial mouth were considered to have fractured at 98 N,

which was the cyclic load in the artificial environment.

Results 52

Between the samples that survived 1,200,000 dynamic loading cycles, the lowest strength

value was observed in test group G (prepared zirconium dioxide implants with crowns;

403.25 N), whereas the highest value was observed in the control group B (one-piece

titanium implants; 7006.77 N).

• Group A: Brånemark® Ti-Implants

- Subgroup A1 (no exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 895.02 834.64 1001.06 784.68 905.13 759.04 662.92 754.70

F= load that led to fracture of the abutment screw

- Subgroup A2 (exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 803.15 702.17 881.63 734.06 931.18 98.00 776.43 793.47

• Group B: Nobel Direct® Ti-Implants (Fig. 27)

- Subgroup B1 (no exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 7006.77 2896.25 4598.23 7004.16 3219.27 7004.60 7002.97 7005.94

- Subgroup B2 (exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 2589.65 2871.62 2161.74 2461.05 3098.84 2247.20 3441.03 3121.34

Results 53

• Group C: Y-TZP BIO-HIP Sigma® Implants

- Subgroup C1 (no exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 1632.32 1224.05 1089.03 1287.01 1138.34 1524.62 1434.13 1369.90 F= load that led to fracture of the implant

- Subgroup C2 (exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 98.00 98.00 1684.75 1476.27 98.00 1941.34 98.00 1349.36

• Group D: Uncoated Y-TZP-A BIO-HIP ® Implants (Fig. 28)

- Subgroup D1 (no exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 838.93 1101.61 804.01 811.64 1021.32 1033.25 1000.76 907.22

- Subgroup D2 (exposure to the artificial mouth with thermocycling)

Sample# 1 2 3 4 5 6 7 8

F (N) 879.84 855.72 870.49 760.68 800.34 974.45 842.86 1044.55

- Subgroup D3 (exposure to the artificial mouth without thermocycling)

Sample# 1 2 3 4 5 6 7 8

F (N) 1479.40 98.00 906.32 937.31 900.05 1097.77 1214.01 1209.03

Results 54

• Group E: Y-TZP-A BIO-HIP Ziunite® Implants

- Subgroup E1 (no exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 743.40 749.60 1000.41 1101.38 954.89 769.68 768.56 710.51

- Subgroup E2 (exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 734.36 754.51 738.18 822.98 98.00 797.71 876.91 978.16

• Group F: Y-TZP-A BIO-HIP Ziunite® Implants

- Subgroup F1 (0.5 mm chamfer preparation, no exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 530.33 576.95 532.97 662.55 618.36 542.22 615.85 548.32

- Subgroup F2 (0.5 mm chamfer preparation + exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 484.54 478.73 734.10 601.68 591.68 569.35 591.59 804.65

• Group G: Y-TZP-A BIO-HIP Ziunite® Implants with crowns

- Subgroup G1 (0.5 mm chamfer preparation + exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 582.35 536.55 568.01 616.99 521.17 638.01 403.25 470.94

- Subgroup G2 (1 mm chamfer preparation + exposure to the artificial mouth)

Sample# 1 2 3 4 5 6 7 8

F (N) 493.29 630.18 576.58 656.57 466.57 497.96 524.96 463.79

Results 55

5.2.2 Statistical evaluation of the data

The mean, median, maximum, minimum and three quartile fracture strength values were

summarized in Table 10 according to the groups. A box plot representation provided a

graphical summary of the data (Fig. 26).

Group Min. 1st Quartile Median Mean 3rd Quartile Max. A1 (Control) 663 757.95 810 825 897.54 1001

A2 (Control) 98 726.08 785 715 822.77 931

B1 (Control) 2896 4253.49 7004 5717 7004.93 7007

B2 (Control) 2162 2407.58 2731 2749 3104.46 3441

C1 (Test) 1089 1202.62 1328 1337 1456.75 1632

C2 (Test) 98 98 724 855 1528.39 1941

D1 (Test) 804 832.10 954 940 1024.30 1102

D2 (Test) 761 832.23 863 879 903.49 1045

D3 (Test) 98 904.75 1018 980 1210.27 1479

E1 (Test) 711 748.05 769 850 966.27 1101

E2 (Test) 98 737.22 776 725 836.46 978

F1 (Test) 530 539.90 563 578 616.47 663

F2 (Test) 479 548.14 592 607 634.78 805

G1 (Test) 403 508.61 552 542 591.01 638

G2 (Test) 464 486.61 511 539 589.98 657

Table 10 Statistical analysis of the fracture strength results in N

In the box plot graph the central box shows the 1st-quartile and the 3rd-quartile and the

line inside the box represents the median value. "Whiskers" (the lines sticking out of the

left and the right end of each box) represent the extremes of the data (they end to the

minimum and maximum value), and very extreme values are shown by themselves

(outliers).

Results 56

Fig. 26 Box plot of the fracture strength test results in N.

Control Subgroup A1: Brånemark® Ti implants (no exposure to the artificial mouth) Control Subgroup A2: Brånemark® Ti implants (exposure to the artificial mouth)

Control Subgroup B1: Nobel Direct® Ti implants (no exposure to the artificial mouth)

Control Subgroup B2: Nobel Direct® Ti implants (exposure to the artificial mouth)

Test Subgroup C1: Y-TZP BIO-HIP Sigma® implants (no exposure to the artificial mouth)

Test Subgroup C2: Y-TZP BIO-HIP Sigma® implants (exposure to the artificial mouth)

Test Subgroup D1: Uncoated Y-TZP-A BIO-HIP® implants (no exposure to the artificial mouth)

Test Subgroup D2: Uncoated Y-TZP-A BIO-HIP® implants (exposure to the artificial mouth)

Test Subgroup D3: Uncoated Y-TZP-A BIO-HIP® implants (exposure, no thermocycling)

Test Subgroup E1: Y-TZP-A BIO-HIP Ziunite® implants (no exposure + no preparation of the implant heads)

Test Subgroup E2: Y-TZP-A BIO-HIP Ziunite® implants (exposure + no preparation of the implant heads)

Test Subgroup F1: Y-TZP-A BIO-HIP Ziunite® implants (no exposure + 0.5 mm chamfer)

Test Subgroup F2: Y-TZP-A BIO-HIP Ziunite® implants (exposure + 0.5 mm chamfer)

Test Subgroup G1: Y-TZP-A BIO-HIP Ziunite® implants with crowns (exposure + 0.5 mm chamfer)

Test Subgroup G2: Y-TZP-A BIO-HIP Ziunite® implants with crowns (exposure + 1 mm chamfer)

Results 57

The results of the multiple pair-wise comparisons of the fracture strength of certain

groups (A1-E1, A2-E2, C1-E1, C2-E2, D2-D3, D1-E1, D2-E2, E1-E2, E1-F1, E2-F2, F1-

F2 and G1-G2) via Wilcoxon rank sum test showed that only the groups C1 vs. E1 and

E1 vs. F1 differed significantly (p<0.05). All the other groups compared demonstrated no

significant difference (p>0.05) (Table 11).

Subgroup p-value Significance

A1 versus E1 1.0000 not significant

A2 versus E2 1.0000 not significant

C1 versus E1 0.0180 significant

C2 versus E2 1.0000 not significant

D2 versus D3 1.0000 not significant

D1 versus E1 1.0000 not significant

D2 versus E2 1.0000 not significant

E1 versus E2 1.0000 not significant

E1 versus F1 0.0163 significant

E2 versus F2 0.1506 not significant

F1 versus F2 1.0000 not significant

G1 versus G2 1.0000 not significant

Table 11 Multiple pair-wise comparisons of the different subgroups via Wilcoxon rank sum test

(significantly different when value p<0.05)

5.2.3 Fracture patterns of the samples In the present study all samples of the control group A fractured at the level of the

abutment screw (Fig. 23), whether in control group B a bending of the titanium one-piece

implants was observed (Fig. 27).

In the test groups most of the samples fractured in a homogenous manner. In groups D, E,

F and G almost all implants fractured at the level of the Technovit® resin (Fig. 28). No

fracture of the zirconia crowns in group G was observed. In group C partial fractures of

the implant heads were also reported (Fig. 29).

Results 58

Fig. 27 Bent head of a Nobel Direct® implant Fig. 28 Fracture of an uncoated Y-TZP-A BIO-HIP®

after the test in the Zwick® machine implant after the test in the Zwick® machine

Fig. 29 Partial fracture of the abutment head of sample No. 5

after the test in the Zwick® machine

Discussion 59

6. Discussion

6.1 Zirconium dioxide one-piece implants as material of interest

The aim of the present study was to evaluate the survival rate and the fracture strength of

different kinds of one-piece zirconium dioxide implants, imitating anterior tooth

replacement, and to compare the results with those of two-piece titanium implants with

an external hexagonal design.

The selection of zirconium dioxide for the fabrication of the implants has been based on

clinical experience with this material in other fields of medicine and dentistry where good

mechanical properties, like a high flexural strength (900 to 1200 MPa), hardness (1200

Vickers) and Weibull modulus (10 to 12), biocompatibility, and low solubility have been

reported (Albrektsson et al. 1985; Christel et al. 1989; Ichikawa et al. 1992; Akagawa et

al. 1993; Akagawa et al. 1998; Piconi and Maccauro 1999). In addition, the white colour

and biotechnical characteristics of zirconia seem to allow the fabrication of high quality

and esthetic reconstructions, which is of greater importance in the areas of the upper

anterior and premolar teeth, especially if the soft tissue situation is not optimal

(Wohlwend et al. 1996; Heydecke et al. 1999).

Initially, the essential tenets for obtaining osseointegration dictated the atraumatic

placement of a titanium screw into bone and a prolonged undisturbed, submerged healing

period. By definition, this required a 2-stage surgical procedure. The external hexagon

design, ad modum Brånemark, originally intended as a coupling and rotational torque

transfer mechanism, consequently evolved by necessity into a prosthetic indexing and

antirotational mechanism (English 1992). With few exceptions, most of the long-term

clinical data on implant performance reported in the literature involve the external

hexagonal, which in each original context of utilization, was used to restore the

completely edentulous arch and performed quite well (Brånemark et al. 1977; Adell et al.

1981). In more complex, partially edentulous and single-tooth applications, the interface

and its connecting screw are exposed to more rigorous load applications (Rangert et al.

1995). The retaining screw is no longer shielded from stress and is subject to lateral

Discussion 60

bending loads, tipping, and elongation, can result in joint opening and screw loosening or

fracture (Jörneus et al. 1992; Sakaguchi and Borgersen 1993; Haack et al. 1995).

Hexagonal screw joint complications, consisting primarily of screw loosening, were

reported as the most common in the literature and ranged from 6% to 48% (Zarb and

Schmitt 1990; Jemt et al. 1992; Jemt and Pettersson 1993; Hemmings et al. 1994; Kallus

and Bessing 1994; Lekholm et al. 1999; Binon 2000; Brägger et al. 2001; Berglundh et

al. 2002; Goodacre et al. 2003; Göthberg et al. 2003). In the present study one implant

from subgroup A2 (two-piece Brånemark® implants) fractured at screw level during the

exposure in the artificial mouth.

Although during the past 10 years controlled torque application and altered screw designs

(i.e. gold-alloy screws) have significantly improved performance (Jörneus et al. 1992;

McGlumphy et al. 1998; Scholander 1999), they have not eliminated the joint problem

entirely. Strub and Gerds showed in an in-vitro study that gold screws and an adequate

preload cannot always guarantee screw joint stability. Once the preload decreases to a

critical level, the external load rapidly erodes the remaining preload. The resulting

vibration, micromovement, and joint interface opening lead to screw loosening and joint

failure (Strub and Gerds 2003). To overcome some of the inherent design limitations of

the external hexagonal connection, a variety of internal alternative connections have been

developed (i.e. cone screw, cone hex, internal octagonal, internal hexagonal, cylinder hex

etc.), in which the screw takes little or no load and provides intimate contact with the

implant walls to resist micromovement, resulting in a stronger interface (Binon 2000).

Möllersten et al. indicated the strength advantage of an internal connection (Möllersten et

al. 1997). Maeda et al. demonstrated that fixtures with internal hex show widely spread

force distribution down to the fixture tip compared with external hex implants (Maeda et

al. 2006).

Additionally some manufacturers have introduced single-component implants,

introducing a new concept for the implant selection process. Parel and Schow evaluated

early clinical results through observation and collection of survival data for 45 Nobel

Direct® implants. The overall success rate was 97.8 % for an observation period from 2.5

to 32 months (Parel and Schow 2005). One-piece implants demonstrate several

advantages. A study by Hermann et al., where 59 implants were placed in edentulous

Discussion 61

mandibular areas of five foxhounds, showed that the biologic width dimension for

transmucosal, one-piece implants, with the rough/smooth border located at the bone crest

level was significantly smaller (p<0.05) compared to two-piece implants with a microgap

(interface) located at or below the crest of the bone. In addition, for one-piece implants,

the highest point of the gingival margin was located significantly more coronally

(p<0.05) compared to two-piece implants (Hermann et al. 2001). These findings, as

evaluated by nondecalcified histology under unloaded conditions in the canine mandible,

suggest that the gingival margin and biologic width dimensions are more similar to

natural teeth around one-piece nonsubmerged implants compared to either two-piece

nonsubmerged or two-piece submerged implants (Hermann et al. 2000). Further studies

have also demonstrated that an intense inflammatory process and thus significantly

greater bone-loss was observed around two-piece implants when compared with one-

piece implants, as the gaps, cavities and hollow spaces, that have been described in two-

piece implants, can be a trap for bacteria, even when good marginal fit of implants

components is present (King et al. 2002; Broggini et al. 2003; Piattelli et al. 2003).

For every day practice using the one-piece design, most often the surgery can be made

flapless, with minimal surgical invasion (Parel and Schow 2005). A flapless approach,

although requiring a greater surgical experience, as the clinician is unable to directly

visualize the path of the osteotomy, has been shown to have inherent benefits in soft

tissue preservation (Becker et al. 2005). It has also been reported to be a safe procedure

in terms of implant success, with percentages equivalent to healed site and delayed

loading protocols (Schwartz-Arad and Chaushu 1998; Campelo and Camara 2002;

Covani et al. 2004). Another benefit of a one-piece implant design is that the implant can

be inserted and immediately restored with a provisional crown. This is of importance in

cases of single-tooth replacement in the esthetic region. It is proposed that implants

immediately restored with a fixed provisional crown enable the enhancement of gingival

contours and generate interdental papillae, thus simulating the natural dentition (Kinsel et

al. 2000). Patient discomfort is also reduced due to the less invasive procedure.

With regard to implantation, one disadvantage is that the implants have to be inserted into

the perfect anatomical position, as only small corrections of the abutment’s inclination

are possible. Without the flexibility of an interchangeable abutment, the initial

Discussion 62

positioning of those implants in the esthetic zone becomes even more critical with a one-

piece design (Parel and Schow 2005).

Considering the above mentioned disadvantages of two-piece implant design, in relation

to the increasing number of patients that seem to prefer tooth-colored, metal-free

materials, which are biocompatible and do not hamper the esthetic appearance of the

entire reconstruction, one-piece zirconium dioxide implants were selected for the present

study.

6.2 Overview of the results In the present study some of the samples have been exposed to the artificial mouth to

simulate 5 years of service before the fracture strength test was performed. The survival

rates of the samples after exposure to the artificial mouth have been 87.5% for subgroups

A2 (Brånemark® Ti implants), D3 (Uncoated ZrO2 implants) and E2 (Ziunite® ZrO2

implants) and 50% for subgroup C2 (Sigma® ZrO2 implants). In the other subgroups a

survival rate of 100% was observed. Then, all test samples (with the exception of those

that fractured in the artificial mouth) were loaded until fracture in a universal testing

machine. The reported mean fracture strength values of the subgroups that were not

exposed to the artificial mouth were: A1 (control) 825 N, B1 (control) 5717 N, C1 (test)

1337 N, D1 (test) 940 N, E1 (test) 850 N and F1 (test) 578 N, and for the subgroups that

were exposed to the artificial mouth were: A2 (control) 715 N, B2 (control) 2749 N, C2

(test) 855 N, D2 (test) 879 N, D3 (test) 980 N, E2 (test) 725 N, F2 (test) 607 N, G1 (test)

542 N and G2 (test) 539 N.

The question, which arises, is whether these values could guarantee the successful

clinical performance and long-term survival of the ceramic implant systems in the oral

cavity.

Discussion 63

6.3 Occlusal forces in the anterior region – Clinical relevance with

the fracture strength values of ceramic materials The average maximum sustainable biting force is approximately 756 N (Gibbs et al.

1986). However, the range of biting forces varies markedly from one area of the mouth to

another and from one individual to another. Higher biting loads are measured on molar

teeth than on other teeth in the dental arch. These forces increase with growth and they

are greater in young adults than in children (Kiliaridis et al. 1993). Although there is a

considerable overlap, biting forces are generally higher in males than in females. This

difference between sexes is due to the fact that men have bigger muscles than women. A

positive correlation between the maximal bite force in the incisor region and the vertical

proportions of the anterior facial morphology has been also reported; subjects with a high

bite force have a relatively short lower anterior height (Garner and Kotwal 1973;

Kiliaridis et al. 1993).

For the incisor region, bite forces range from 60 N to 360 N (Bates et al. 1976; Kiliaridis

et al. 1993). These values show variations between different authors.

Helkimo et al. reported average values of 147 N for the canines and 137 N for the

incisors. The mean values were higher for males than for females. In males the maximal

bite forces measured were 176 N in the incisor region and in females 108 N (Helkimo et

al. 1977).

Sonnenburg et al. found that the average load values for anterior teeth are between 215 N

and 360 N for men and 115 N and 269 N for women (Sonnenburg et al. 1978).

Kalipcilar and Kedici (1993) reported maximal biting forces of 121.6 N for the canines

and of 94.1 N for the incisors (Kalipcilar and Kedici 1993).

When comparing with the maximal biting forces exerted by the stomatognathic system, it

is obvious that average functional chewing forces are smaller. Normally, the energy of

the bite is absorbed by the food bolus during mastication, as well as by the tooth or

implant, periodontal ligament (as far as teeth are concerned), and bone. On the other hand

the magnitude of chewing forces placed on implant-supported fixed reconstructions has

been shown to be considerably greater than the one placed on teeth in dentate patients.

This may depend on the lack of the periodontal proprioceptive feedback in the implant-

Discussion 64

supported reconstructions (Kim et al. 2005). It should be also mentioned that the incisors

are not usually involved in the process of chewing, but most normally act as cutters

(Anderson 1956; Eichner 1963; Bates et al. 1976; De Boever et al. 1978).

Therefore, it is more realistic, for purposes of experimental design, to consider higher

loading forces than the functional forces that arise during chewing or swallowing. In this

study during the exposure to the artificial mouth, a cycle loading force of 98 N was used

in order to approach a clinical situation comparable to the anterior area of the dental arch.

In-vitro studies have shown that the minimum initial fracture strength of ceramic

materials used for front-teeth restorations, without being previously exposed to the

artificial mouth, should be 400 N, as these materials should be able to withstand the

maximum of occlusal forces applied on natural teeth (Schwickerath 1988). In another

study cylindrical ceramic test specimens were subject to fatigue loading in bending tests.

After 1000 loadings the strength was found to be approximately 60% of the value

obtained in a static bending test. A marked reduction in strength under fatigue loading in

bending tests was observed leading to the conclusion that in determining the maximum

load values for the clinical appraisal of ceramic materials the fatigue limit should also be

taken into consideration (Schwickerath 1986). Körber and Ludwig reported that for

ceramic materials a fracture strength, which is the 60% of the initial strength, is satisfying

for the clinical use of these materials (Körber and Ludwig 1983). Therefore, samples

exposed to the artificial oral environment should exhibit minimum fracture strength

values of 200 N, when imitating anterior teeth restorations, and 300 N, when imitating

posterior teeth restorations.

In our investigation, the observed mean fracture strength values of the subgroups that

were not exposed to the artificial mouth were greater than 400 N (reported mean values >

578 N). For the samples that were exposed to the artificial mouth the mean fracture

strength values were greater than 200 N (reported mean values > 539 N) and within the

limits of clinical acceptance, when compared with the physiologic forces developed in

the oral cavity, as suggested by Körber and Ludwig (Körber and Ludwig 1983). These

values are also greater than 300 N, which is the minimum fracture strength value

proposed for restorations of the posterior region.

Discussion 65

6.4 The clinical relevance of fracture strength tests

Fracture strength tests of ceramic materials are of substantial importance in deciding the

efficiency, longevity and subsequent success of the reconstructions. Ideally, in-vitro load-

to-failure tests should simulate clinical failure, in order to evaluate new materials or

designs and recommend them for clinical use (Ritter 1995a).

The traditional fracture strength tests (biaxial flexural tests or 3-, 4-point bending tests)

cannot be used to predict the performance of a restoration in the mouth, because they do

not simulate a clinical mode of failure (Kelly 1995; Ritter 1995b). Also, traditional

fracture strength tests can fail from their edges during processing, raising the possibility

that premature failures occur from the uncharacteristic processing flaws (Kelly 1995).

Therefore, in order for testing to be relevant, test specimens must have the same type and

distribution of flaws as the target structure when placed into service (Kelly 1995; Ritter

1995a, b).

Kelly questioned the clinical validity of such tests (Kelly 1999). He suggested that

significant differences were found between the failure behaviour created during

traditional load-to-failure tests and that observed to have occurred during clinical failure

of all ceramic restorations. He mentioned that the elementary beam theory couldn’t be

used to examine a cemented full coverage restoration or to predict its clinical behaviour.

Homogenous all-ceramic restorations consist of a layer of ceramic, a layer of cement and

are supported by a thickness of dentin. This structure of different materials having

different modulus of elasticity is not well represented by simple bar-shaped specimens,

such as those used in 3-point or 4-point bending tests. Another problem of laboratory

testing is that extremely high loads are required compared with those during mastication,

swallowing or clenching movements. Furthermore, water showed that it acts chemically

at crack tips and decreases the strength of glass and ceramics. He concluded that

traditional laboratory tests fail to 1) create appropriate stress states, 2) cause failure from

clinically relevant flaws, or 3) create crack systems modelled on clinical failure. No

conclusion regarding longevity of the restoration can be drawn from fracture strength

results alone.

Discussion 66

For the strength test to accurately reflect the variability and time dependency of a ceramic

component in service, the test environment must be the same as the service environment,

and the strength-controlling flaw population must be the same as the responsible for

failure in service. Therefore, it is generally recommended that the test specimens and

mode of loading be chosen to closely simulate the actual components in service (Ritter

1995a, b; Kelly 1999).

In the present study, the specimens were tested in the Zwick universal testing machine,

until fracture occurred. This might not closely simulate the service environment, but it

gives a hint on the stability of the specimens in clinical use. A more elaborate test setting

would be of course desirable. However, our specimens were set at an angle of 130o to

their long axis, like in the artificial mouth, which imitates the clinical testing. The

angulation is in agreement with the intraoral orthognathic interincisal angle of 130o (Reitz

et al. 1973).

6.5 Clinical relevance and influence of the artificial mouth on the

survival rate and fracture strength of ZrO2 implants

Before performing in-vivo studies or applying new dental materials for clinical use, in-

vitro tests are recommended in order to prove their applicability and performance. In-

vitro tests can be performed in a short period of time and have the advantages of

reproducibility and the possibility of standardising the test parameters (Krejci and Lutz

1990; Kern et al. 1999). However, each in-vitro test represents only one approach to a

clinical situation. The most closely a test simulates the clinical situation, the more

clinically relevant are the results (Krejci et al. 1990).

It has been shown that ceramic materials accumulate damage during cycling loading and

thermocycling. The accumulated subsurface damage weakens the ceramic and can cause

clinical failures (Kelly 1999). Intraorally, occlusal forces create dynamic repetitive

loading. Therefore, instead of monotonic static loading, it is more clinically relevant to

test the specimens under physiologic fatigue load in a chewing simulator, which will

allow evaluation of the dental restorative systems under clinically relevant conditions

(Schwickerath 1986; Krejci and Lutz 1990; Krejci et al. 1990; Geis-Gerstorfer and Fäßler

Discussion 67

1999; Kelly 1999; Kern et al. 1999). Krejci et al. suggested that the chewing machine

fulfils the parameters concerning chewing motion and thermal changes reported in the

literature (Krejci et al. 1990). For the above reasons, in the present study a dual-axis

chewing simulator was used for the artificial ageing.

The parameters used for the chewing machine were limited to the physiological values

found in the literature. A force of 49 N was applied during the dynamic loading in several

studies (Krejci and Lutz 1990; Krejci et al. 1990; Kern et al. 1999). This reflects mean

values of chewing forces in the anterior dentition according to the literature. However,

clinical studies showed that biting forces could easily exceed the 49 N loading forces

(Gibbs et al. 1986). In this study, a cycle loading force of 98 N was used in order to

approach a clinical situation comparable to the anterior area of the dental arch.

Adding moisture and controlled temperature to the environment was found to be

important when measuring the fracture or fatigue strength of dental ceramics. Exposure

to water was found to affect the mechanical properties of all-ceramic restorations (Kelly

et al. 1989; Drummond et al. 1991; Kelly 1995). Furthermore, temperature changes also

lead to slow flaw propagation (Kelly 1999). Although some authors used a median

temperature of 37oC for testing dental materials in the artificial oral environment

(Kappert and Altvater 1991), most of the authors used temperatures varying from 5.0oC

to 55oC for their tests (Beschnidt and Strub 1999; Kern et al. 1999). This is in accordance

with Palmer et al. (1992) who reported minimum and maximum temperatures intraorally

between 1.0oC and 58.5oC and suggested a range of 0oC to 67oC for the thermocycling

tests of dental materials (Palmer et al. 1992). In this study temperatures varying from

5.0oC to 55oC were used simultaneously with the cyclic loading. In subgroup D3 the

absence of water thermocycling during the exposure of the implants to the artificial

mouth may have led to the higher mean fracture strength value of these implants (980 N)

in comparison to the implants of subgroup D2, which were exposed to the artificial

mouth with water thermocycling (879 N). However, among the groups compared (D2-

D3) no statistically significant difference was reported (p>0.05).

Clinical studies showed that humans have an average of 250,000 masticatory cycles per

year (DeLong and Douglas 1983; Sakaguchi et al. 1986). Therefore, to simulate a service

time of 5 years, about 1,200,000 masticatory cycles have to be performed in the chewing

Discussion 68

simulator (Krejci and Lutz 1990; Kern et al. 1999). A chewing cycle in the artificial oral

environment is designed to correspond as closely as possible to physiological conditions.

The magnitude, duration and frequency of the force applied in the artificial mouth are

comparable to values reported in the literature (Bates et al. 1975).

In this study, balls made of Steatite (a ceramic enamel analogue) were used as antagonists

for the test samples in the chewing machine. It has been shown that antagonists made out

of enamel are not suitable for standardised wear tests, because they differ in morphology,

microstructure and composition (Wassell et al. 1994b). Metals and composites are also

not acceptable as antagonists (Krejci et al. 1990). The advantages of Steatite antagonist

are that it has a Vickers hardness that is very similar to enamel, and is accurate,

reproducible and cost effective (Wassell et al. 1994a, b).

In the present investigation the survival rates of the samples after exposure to the

artificial mouth have been 87.5% for subgroups A2, D3 and E2 and 50% for subgroup

C2. In the other subgroups a survival rate of 100% was reported. The sample of subgroup

A2 fractured at the abutment screw level, which could be attributed to the deficiency of

external hex connection, as described above. An explanation for the lower survival rate of

subgroup C2 compared with the other subgroups could be the quality and microstructural

design of the zirconia used for the fabrication of the Sigma® implants, as all zirconia

materials are not the same (Cales et al. 1994). However, when the mean fracture strength

values of the Sigma® implants of subgroup C1 (1337 N) and of the Ziunite® implants of

subgroup E1 (850 N) were compared a statistically significant difference was observed

(p<0.05). These results demonstrate a higher stability and strength of the Sigma®

implants when they are not exposed to the artificial mouth, which contradicts the high

failure rate of these implants in the artificial mouth. This could be attributed to a material-

related higher metastability rate (t→m phase transformation) in the Sigma® implants in

comparison to the Ziunite® implants leading to material degradation during the exposure

to the artificial mouth. For the 87.5% survival rate of subgroups D3 and E2 no possible

explanation could be given.

The mean fracture strength values of the samples for most of the subgroups exposed to

the artificial mouth were lower when compared to the mean values of the samples that

were not exposed. That was expected as the exposure to the artificial mouth results in

Discussion 69

aging of the examined materials. However, the question whether the exposure of the

samples to the artificial mouth has a statistically significant influence on the fracture

strength of the implants or not was negatively answered by our results. The difference of

the groups compared (E1 vs. E2 and F1 vs. F2) was not found to be statistically

significant (p>0.05).

In his thesis Klaus investigated the fracture strength of zirconium dioxide implants,

which were fabricated ad modum ReImplant® (two piece root-analogue implants) and

were restored with IPS Empress® 1 ceramic or Procera® aluminium oxide crowns, and he

compared them with titanium implants, which were restored with metal-ceramic crowns.

One subgroup (8 implants) from each group (16 implants) was exposed to 1.2 million

thermomechanical loading cycles in the artificial mouth. Similarly to the present study,

he found that the exposure to the artificial mouth had not a statistically significant

influence on the fracture strength values of the implants. The survival rate of all samples

after the exposure was 100%. Klaus concluded that the systems examined could

withstand 5 years of clinical service under normal functioning conditions (Klaus 2002).

However the applied by Klaus weight per sample was 30 N, whereas the applied load in

our investigation was 98 N.

In a similar study, Finke examined the fracture strength of zirconium dioxide implants,

which were also fabricated ad modum ReImplant® (two-piece cylindrical implants) and

were restored with crowns made out of IPS Empress® 2 ceramic or zirconium dioxide

(veneered with Triceram®). He compared them with titanium implants, which were

restored with metal-ceramic crowns. One subgroup (8 implants) from each group (16

implants) was exposed to 1.2 million thermomechanical loading cycles in the artificial

mouth with an applied weight of 30 N per sample. The survival rate of the implants

restored with the IPS Empress® 2 crowns was 100%, of those restored with the zirconium

dioxide crowns was 87.5%, and of the control group was 25%. The reason for the low

survival rate in the control group (titanium implants with metal-ceramic crowns) was the

fracture of the abutment screw. The explanation given by Finke for the low survival rates

of the samples in the control group was that the simultaneous use of cement with the

purpose of a more stable abutment-implant connection has led to a misfit of the implant

components resulting in the fracture of the abutment screws. Similarly to the present

Discussion 70

study, Finke reported no statistically significant difference of the mean fracture strength

values between the groups that were exposed and those that were not exposed to the

artificial mouth (Finke 2004). However, the mean fracture strength values reported by

Finke, were for the first test group (ZrO2 implants-Empress®2 crowns) with and without

exposure to the artificial mouth 280.7 N and 281.8 N, for the second test group (ZrO2

implants -Triceram® crowns) 275.7 N and 325.1 N and for the control group (Ti-metal-

ceramic crowns) 165.7 N and 595.2 N respectively (Finke 2004). These values are

considered to be very low and they would not allow the clinical use of this implant

system.

6.6 Surface and heat treatments of zirconia devices (abutments,

crowns)

The influence of the implant abutment preparation and surface treatment of the all-

ceramic crowns is a subject of discussion. Several authors have investigated the effect of

sandblasting, wet and dry grinding on the mechanical properties of Y-TZP ceramics. As

already explained the tetragonal grains of zirconia, which are normally stable at high

temperatures, can be retained at room temperature by adding metal oxides, such as yttria

(Y2O3) (Piconi and Maccauro 1999). Nevertheless, the tetragonal grains may transform

into monoclinic as a result of externally applied stresses exerted by grinding and

sandblasting, for example. The tetragonal to monoclinic (t→m) phase transformation

exhibits a 4% volume expansion that creates compressive stresses at the crack tip. These

stresses must be overcome by the crack in order to propagate, explaining the greater

fracture toughness of zirconia compared to conventional dental ceramics (Piconi and

Maccauro 1999).

However, the influence of grinding on the flexural strength of zirconia ceramics is

contradictory and related to the volume percentage of transformed zirconia, which in turn

depends on the metastability of the t→m phase transformation, the grinding severity and

the locally developed temperatures (Gupta 1980; Green 1983; Swain 1985; Kosmac et al.

1999). Regarding ZrO2, grinding has been recommended to create a surface region of

compressive stresses, which increases the mean flexural strength of zirconia ceramics. On

Discussion 71

the other hand severe grinding introduces deep surface flaws which may become strength

determining if their length largely exceeds the depth of the grinding induced surface

compressive layer (Gupta 1980; Green 1983; Denry and Holloway 2006).

Luthardt et al. demonstrated a 50% reduction in flexural strength of Y-TZP zirconia

ceramics machined under conditions simulating the inner surface grinding of crowns and

FPDs when compared with the controls (Luthardt et al. 2002).

Swain and Hannink showed that hand grinding is more effective than lapper-machine

grinding in inducing the t→m transformation. In the case of machining grinding the local

development of temperatures exceeded the m→t transformation temperature, causing the

reverse m→t transformation. In this instance, the deep defects introduced by grinding are

no longer counteracted by the transformation-induced compressive stresses and act as

stress concentrators, lowering the mean flexural strength of the ceramic (Swain and

Hannink 1989).

Kosmač et al. investigated the influence of sandblasting, wet and dry grinding in dry

pressed and sintered 3mol% Y2O3 TZP. They showed that sandblasting was more

effective than grinding in inducing the t→m transformation and therefore, increasing the

mean flexural strength of the ceramic. Sandblasting was described as a process able to

induce transformation without developing high temperatures or creating severe surface

damage, and therefore, strengthening the material (Kosmac et al. 1999).

In another study Guazzato et al. also showed that the flexural strength of Y-TZP ceramics

were increased by sandblasting and wet-grinding, while lower mean strength values were

measured when the same procedures were followed by heat treatment (Guazzato et al.

2005). They concluded that a greater amount of monoclinic phase and therefore a greater

flexural strength could be obtained for sandblasted and ground specimens, and may be

clinically of advantage. On the other hand, the gain in flexural strength achieved through

the compressive stresses created by the phase transformation may be lost as a result of the

exposure of the transformed surface to an aqueous environment. It has been shown that

the formation of monoclinic phase on the surface of the material after exposure to an

aqueous environment is accompanied with microcracking. This is due to the absorption

of water to Zr-O bond and the formation of surface hydroxyls (Sato and Shimada 1985;

De Aza et al. 2002; Guazzato et al. 2005).

Discussion 72

Contradictory were the conclusions of Fischer et al. who examined the material

properties of zirconia abutments that were subjected to preparation procedures with and

without water coolant, comparing them to untreated samples. They demonstrated that the

polymorphism of the material was only minimally influenced by the machining process

(with and without water cooling), whereas the strength of the prepared abutments proved

to be at the same level as that of the untreated blanks (Fischer et al. 1999).

In the present study the results confirm the findings of most of the above mentioned

studies that wet-grinding procedures increase the flexural strength of Y-TZP, resulting in

a more stable but weaker material. All of the implants of subgroups F2, G1 and G2,

which were submitted to grinding procedures, survived in the artificial mouth (greater

stability compared to the implants of group E2), but demonstrated mean fracture strength

values lower than the implants of subgroup E2 (unprepared implants). From our results, it

was obvious that the preparation of the abutments had a negative influence in the fracture

strength values of the implants. The temperature changes that happen during the

preparation of the abutment could lead to aging of the material, resulting in lower fracture

strength values (Guazzato et al. 2005). When the mean values of subgroup E1

(unprepared implants) were compared with those of subgroup F1 (prepared implants) the

difference was found to be statistically significant (p<0.05). However, the difference

between subgroups G1 and G2 when compared (different width of chamfer preparations

and restoration with all-ceramic crowns) was not reported to be statistically significant

(p>0.05).

6.7 Influence of the surface topography and the design of the

implants on the fracture strength values In the present study the implants of three groups (E, F and G) had a special surface

preparation. The implant body was covered with slurry containing zirconia powder and a

pore-former, so that a porous surface could be achieved. The rationale was to achieve

improved stability and attachment in cortical bone as well as better mechanical locking in

medullar bone with this roughened surface of the implant. The same surface preparation

of the implant bodies was examined by Sennerby et al. (Sennerby et al. 2005). The

Discussion 73

authors showed a strong bone tissue response to surface-modified zirconia implants after

6 weeks of healing in rabbit bone. The modified zirconia implants showed a resistance to

torque forces similar to that of oxidized implants and a four- to fivefold increase

compared with machined zirconia implants. They concluded that the surface modified

zirconia implants can reach firm stability in bone (Sennerby et al. 2005). However,

regarding the results of our investigation no significant difference of the mean fractures

strength values has been found between the implants with and without a special surface

topography (D1 vs. E1 and D2 vs. E2) (p>0.05). It could be concluded that the examined

surface topography does not influence the mechanical stability of the above implants.

As far as the design of the implants is concerned (one-piece implants compared to two-

piece implants), no statistically significant difference was reported between the groups

compared in our study (A1 vs. E1 and A2 vs. E2) (p>0.05). A possible explanation could

be the controlled torque (32 Ncm) applied to the abutment screws of the Brånemark®

implants, which is recommended by the manufacturer as it has been proven to improve

the performance of these implants.

The mean fracture strength values of two-piece implants in the test groups reported by

Finke in his thesis were relatively low when compared with the mean fracture strength

values of subgroups G1 and G2 of the present study (prepared implants exposed to the

artificial mouth-different preparation forms), which were 542 N and 539 N respectively.

This could be related to the different design of the implants used in our study (one-piece

implants), when compared to the two-piece cylindrical implants used by Finke (2004).

One-piece implants show lower shearing forces at the implants neck and that could be a

possible reason for the higher mean fracture strength values of the implants in our study.

However, the different study designs (i.e. different kind of implants, the applied load in

the artificial mouth) do not allow a direct comparison of the mean fracture strength values

reported in both investigations.

6.8 Patterns of fracture of the samples

In the present study the patterns of fracture of the test samples were similar. Almost all

implants fractured at the implant neck at the level of the Technovit® resin. This results

Discussion 74

from the fact that the elements exposed to maximum stress are located around the neck of

the implant (Himmlová et al. 2004). Only in group C partial fractures of the implant

heads were reported. This could be attributed to the fact that the implant heads of the

Sigma® implants have a lower length in comparison to the implants heads of the other

ceramic implants. Consequently, there is a lower lever effect during the loading of the

Sigma® implants.

In a number of finite element studies it has been found that the peak stress in the bone

supporting a dental implant appears in the crestal region close to the level where the

implant starts to attach to the bone (Kitoh et al. 1988; Mailath et al. 1989; Meijer et al.

1993).

Himmlová et al. showed in their simulation study that the implant diameter was more

important for improved stress distribution than implant length. They suggested that the

wider area in the cervical portion of the implant may better dissipate the masticatory

forces (Himmlová et al. 2004).

Hansson et al. also demonstrated in a finite element analysis that if a dental implant is

provided with retention elements all the way up to the crest, as opposed to an implant

with a smooth neck, the axial load, which the implant can bear, will be increased to a

substantial degree. They also suggested that keeping the major diameter possible in each

case and an increased wall thickness of an implant results in an increase in the axial load

the implant can resist (Hansson 1999).

The above-mentioned conclusions could lead to an improvement of the implants design

used in the present study, in which all of the implants had a smooth neck.

One interesting observation was that the preparation of the implant heads reduces the

fracture strength values of the implants (Groups F and G). Although, the preparation

occurred approximately 2-3 mm above the Technovit the specimens fractured at a lower

force still at the Technovit level compared to the unprepared specimens. This may

indicate that this is a remote effect of the preparation, which needs to be investigated.

Which effect leads to the reduced fracture strength at the Technovit level? Is it a thermal

influence that leads to degradation of the material? Generally, it can be ruled out that

flaws are introduced at the Technovit level, which may exceed the thickness of a

toughening monoclinic zirconia layer.

Conclusions 75

7. Conclusions

Within the limits of this in-vitro study it can be concluded that:

1. Y-TZP-A BIO-HIP®can be an alternative material to titanium for dental implant

fabrication, since the survival rates obtained and the mean fracture strength values

were within the limits of clinical acceptance.

2. The performance of one-piece implants made out of Y-TZP BIO-HIP® is not as

good as the performance of the implants fabricated out of Y-TZP-A BIO-HIP®.

They demonstrated a survival rate of only 50% after exposure to the artificial

mouth.

3. The exposure of the implants to the artificial mouth has not a statistically

significant influence on the mean fracture strength values of the implants.

4. The preparation of the abutments has a negative influence on the fracture strength

values of the implants.

5. Although first experimental data are encouraging, long-term clinical data are

necessary before one-piece zirconia implants can be recommended for daily

practice.

Summary 76

8. Summary

The aim of the present study was to evaluate the survival rate and the fracture strength of

zirconium dioxide implants, imitating anterior tooth replacement, and to compare the results with

those of titanium implants before and after exposure to the artificial mouth.

A total of 120 ceramic and titanium implants were used for the experiment. The Ti-implants were

divided into two control groups A (two-piece implants) and B (one-piece implants) including

sixteen titanium implants each. The ceramic implants (one-piece implants) were divided into five

groups C (Y-TZP BIO-HIP® implants), D (Y-TZP-A BIO-HIP® implants), E (Y-TZP-A BIO-

HIP® implants with special surface topography), F (Y-TZP-A BIO-HIP® implants with special

surface topography + preparation of the implant heads) and G (Y-TZP-A BIO-HIP® implants

with special surface topography + restoration with ZrO2 crowns, different width of the

preparation in each subgroup), using sixteen samples of each, with the exception of group D,

which included twenty-four samples. The test and the control groups were divided into 15

subgroups of 8 samples each. One subgroup from groups A (A2), B (B2), C (C2), E (E2) and F

(F2), two subgroups from group D (D2, D3) and both subgroups from group G (G1, G2) were

exposed to 1.2 million thermomechanical loading cycles in the artificial mouth in order to

simulate 5 years of clinical service. Fracture of the implants was considered as failure. Then, all

test specimens (exposed and not exposed to the artificial mouth) were loaded until fracture

occurred in a universal testing machine.

After exposure to the artificial mouth the survival rates reported were as follows: A2 87.5%, C2

50%, D3 (no thermocycling) 87.5% and E2 87.5%. All the other subgroups demonstrated a

survival rate of 100%.

The observed mean fracture strength values of the subgroups that were not exposed to the

artificial mouth were: A1 825 N, B1 5717 N, C1 1337 N, D1 940 N, E1 850 N and F1 578 N. For

the samples that were exposed to the artificial mouth the mean fracture strength values were the

following: A2 715 N, B2 2749 N, C2 855 N, D2 879 N, D3 980 N, E2 725 N, F2 607 N, G1 542

N and G2 539 N. Among the different groups no statistically significant differences were found,

with the exception of subgroup C1 when compared to subgroup E1 and of subgroup E1 when

compared to subgroup F1 (p<0.05). The results in the present study showed that the preparation

of the abutments had a negative influence in the fracture strength values of the implant.

All mean values obtained were within the limits of clinical acceptance, indicating the use of one-

piece zirconia implants may be clinically acceptable. However, long-term clinical data are

necessary before one-piece zirconia implants can be recommended for daily practice.

Zusammenfassung 77

9. Zusammenfassung Ziel dieser Untersuchung war es, die Überlebensrate und Bruchfestigkeit verschiedener

Zirkoniumdioxid-Implantate mit und ohne Alterung in einer Kaumaschine zu erfassen und mit

dem Verhalten von Titanimplantaten unter gleichen Bedingungen zu vergleichen.

In der Studie wurden 120 keramische und Titanium Implantate verwendet. Es wurden zwei

Kontrollgruppen von Titan-Implantaten, Gruppe A (zweiteilige Ti-Implantate) und B (einteilige

Ti-Implantate), zu je 16 Einheiten gebildet. Die keramischen Implantate wurden in 5 Testgruppen

eingeteilt, Gruppe C (Y-TZP BIO-HIP®Implantate), D (Y-TZP-A BIO-HIP®Implantate), E (Y-

TZP-A BIO-HIP® Implantate mit Oberflächenbeschichtung), F (Y-TZP-A BIO-HIP® Implantate

mit Oberflächenbeschichtung + 0.5 mm Hohlkehlpräparation der Implantatpfosten) und G (Y-

TZP-A BIO-HIP® Implantate mit Oberflächenbeschichtung + Versorgung der Implantate mit

ZrO2-Kronen (Procera®), 0.5 mm und 1 mm Hohlkehlpräparation der Implantatpfosten in

Untergruppen G1 und G2). Jede Gruppe umfasste 16 Einheiten, mit Ausnahme der Gruppe D, die

von 24 Implantaten gebildet wurde. Die Kontroll- und Testgruppen wurden in 15 Untergruppen

eingeteilt, zu je 8 Implantaten. Die Untergruppen A2, B2, C2, D2, D3, E2, F2, G1 und G2 wurden

einer Kausimulation von 1,2 Mio. Zyklen ausgesetzt, was in etwa einem klinischen

Funktionszeitraum von 5 Jahren entspricht. Abschließend wurden alle Prüfkörper einem

statischen Bruchbelastungstest unterworfen.

Nach der Exposition im künstlichen Mund wurden die folgenden Überlebensraten beobachtet: A2

87,5%, C2 50%, D3 (ohne Thermolastwechsel) 87,5% and E2 87,5%. Die Überlebensrate der

anderen Gruppen war 100%.

Die mittleren Bruchbelastungswerte betrugen für die Gruppen, die nicht der Kaumaschine

ausgesetzt wurden: A1 825 N, B1 5717 N, C1 1337 N, D1 940 N, E1 850 N und F1 578 N. Die

Werte für die kaumaschinell gealterten Gruppen betrugen: A2 715 N, B2 2749 N, C2 855 N, D2

879 N, D3 980 N, E2 725 N, F2 607 N, G1 542 N und G2 539 N. Der Unterschied zwischen den

Untergruppen C1 und E1 und E1 und F1 war statistisch signifikant (p<0,05)(Wilcoxon

Rangsummen Test). Es konnte gezeigt werden, dass die Präparation der Implantatpfosten einen

negativen Einfluss auf die Bruchfestigkeit der entsprechenden Implantate hat.

Zusammenfassend lässt sich feststellen, dass die Bruchfestigkeit der getesteten keramischen

Implantate vergleichbar mit der Stabilität der Verbindung Pfosten-Implantat bei zweiteiligen

Titanimplantaten ist. Eine klinische Anwendung scheint gerechtfertigt zu sein. Klinische

Untersuchungen müssen jedoch zeigen, ob die keramischen Implantate die anerkannten

Erfolgskriterien der Implantologie erfüllen können.

References 78

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Curriculum vitae 100

11. Curriculum vitae Personal Data: Name: Marina Andreiotelli

Date of birth: March 29th, 1980

Place of birth: Athens, Greece

Parents: Miltiadis Andreiotellis

Eleni Andreiotelli

Sister: Panagiota Andreiotelli

Marital status: Single

Nationality: Greek

Education: 1991-1994 Secondary school, Athens

1994-1997 High school, Athens

1997-2002 University of Athens,

Faculty of Dentistry, Greece

2003-2006 Postgraduate Student, Department

of Prosthodontics, Albert-Ludwig

University, Freiburg, Germany

Program director: Dr. S. Smeekens,

Associate Professor

Work experience: 03/2003- 10/2003 Dentist in the prosthodontic department

of the military hospital N.I.M.I.T.S,

Athens, Greece

Scientific Work: 7 oral presentations

29 participations in scientific congresses

2 publications in Greek scientific journals

Acknowledgements 101

12. Acknowledgements I would like to express my most sincere gratitude to Prof. Dr. R. J. Kohal, Department of

Prosthodontics, Albert-Ludwig University, Freiburg, Germany for offering me the

possibility to conduct research under his supervision, as well as for his revision of this

manuscript.

I would like to express my gratitude to the Alexander S. Onassis Public Benefit

Foundation, Athens, Greece for supporting me financially through a two-year

scholarship.

In addition, I would like to address my acknowledgments to the following persons

without whose help this thesis could not have been completed:

PD Dr. T. Auschill, Department of the Operative Dentistry and Periodontology, Albert-

Ludwig University, Freiburg, Germany, for the revision of the manuscript.

Prof. Dr. J. R. Strub, Chairman of the Department of Prosthodontics, School of Dentistry,

Albert-Ludwig University, Freiburg, Germany, for offering me the opportunity to work in

his department.

Dr. W. Att for his guidance at the start of this research project, his useful advice and

support in operating the artificial oral environment.

Dr. M. Krah, Division of Dental Materials, Albert-Ludwig University, Freiburg,

Germany for his help in carrying out the fracture strength tests.

Mr. T. Gerds, Institute of Medical Biometry and Medical Informatics, Albert-Ludwig

University, Freiburg, Germany for performing the statistical analysis of the data.

I would add a special thank to my dear parents for their continuous moral and financial

support during all circumstances, always respecting and encouraging my wishes.

Acknowledgements 102

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