14089. Direct Ceramic Machining of. Ceramic Dental Restorations. A dissertation
submitted to the. SWISS FEDERAL INSTITUTE OF TECHNOLOGY ZURICH.
Research Collection
Doctoral Thesis
Direct ceramic machining of ceramic dental restorations Author(s): Filser, Frank Thomas Publication Date: 2001 Permanent Link: https://doi.org/10.3929/ethz-a-004183626
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ETH Library
Diss. ETH No. 14089
Direct Ceramic Machining of Ceramic Dental Restorations
A dissertation submitted to the SWISS FEDERAL INSTITUTE OF TECHNOLOGY ZURICH for the degree of Doctor of Technical Science
presented by
Frank Thomas Filser Dipl.-Ing. TU Kaiserslautern born on April 24, 1966 German Accepted on the recommendation of Prof. Dr. Ludwig J. Gauckler, examiner Dr. T. Graule, co-examiner Prof. Dr. Heinz Lüthy, co-examiner
Zurich, 2001
Acknowledgements
Acknowledgements I am grateful to Prof. Gauckler who gave to me the opportunity to realize this thesis in his laboratory. He encouraged, supported and motivated me with much kindness throughout this work. In particular he showed me the interesting sides of materials science and those of the highly interdisciplinary integrating project work. Special thanks further goes to Peter Kocher for his advice, help, and his technical support. We had a lot of fun together throughtout this time, starting with the fabrication of the first dental bridge to the design and the fabrication of the prototype machines. He showed to me the bernian and the italian life style, and the best places to eat in Switzerland. I also recall our time in the United States. I thank the "DCM team" consisting of Prof. Gauckler, Peter Kocher, Prof. Lüthy and Prof. Schärer for their believe in the DCM system, their contribution towards its success and going along the pathway towards its realization together. I benefited from the discussions with Prof. Lüthy on the dental materials and from those with Prof. Schärer on the dentistry, both showing great expertise and never-ending constructive criticism. Dr. Graule and Prof. Lüthy are gratefully acknowledged for being co-examiner with all the affiliated work. I thank Dr. Graule for his persisting interest in this work from the beginning onwards. I am grateful to Prof. Smith for being chairman. I thank Prof. Bayer for all his advice on ceramics and on the English language. He showed experience, knowledge, and infinite patience. He also did the proof reading of the manuscript. Irène Urbánek I thank for doing most of the administrative work and for her support. Most of the administrative things run much easier when asking her.
Acknowledgements
Furthermore, I am in debt to many present and former colleagues at the chair of Nonmetallic Inorganic Materials. In particular, I enjoyed the time with Stefan Köbel who fruitfully discussed any ideas with me, and with whom I had lot of fun snowboarding in Canada and in Switzerland, with Anja Bieberle and Markus Hütter whom I skied in the swiss mountains and visited the highlands, with Julia Will for her humor and frankness and making me sing. I acknowledge my students Brandon Bürgler, Christian Ceppi, Daniela deStefano, Michael Eglin, Reto Haggenmüller, Jost Lussi, Katja Michel, for their valuable contributions to this work. I had much fun "competing" with Urs Schönholzer in fabricating zirconia ETH medals. I thank Nicholas Grundy for proof reading the manuscript. I also wish to thank the University’s group at the Clinic of Fixed and Removable Prosthodontics and Dental Materials Science in Zurich, for many helpful discussions on dentistry and dental technique. My thanks goes especially to Dr. Aurel Fehér, who helped and adviced me in all dental questions, took responsibility for the clinical studies, did most of the dental work and who also treated me with a DCM - crown. Further I thank Madeleine Schumacher for her skilled artistic work on our frameworks and for her expertise, and Olivier Loeffel for his technical support. Financial support from the Swiss Priority Program on Materials research from the board of Swiss Federal Institutes of Technology is gratefully acknowledged. I thank Dr. Reinhard for his advice and his administrative support. Finally, my highest appreciation is addressed to my family for giving me moral support.
I
Table of Contents
Table of Contents
I.
II.
III.
The Direct Ceramic Machining System (DCM) . . . . . . . . . . . . . 1 I.1.
Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1
I.2.
Zusammenfassung . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3
Synopsis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 II.1.
Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6
II.2.
Goal and Approach . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9
II.3.
Outline of the Thesis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11
II.4.
References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12
State of the Art: Fabrication of All-Ceramic Restorations . . . 15 III.1.
Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16
III.2.
Hot Pressing - Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 17
III.3.
Casting - Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 19
III.4.
Slip Casting - Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 20
III.5.
Compaction of Powder . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 21
III.6.
Hard Machining - Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22
III.7.
Soft Machining . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 25
III.8.
Other Approaches . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 27
III.9.
Discussion and Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 29
III.10. References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 31
Table of Contents
IV.
II
All-Ceramic Teeth Restorations: The Direct Ceramic Machining (DCM)-Process . . . . . . . . . . . 39 IV.1. Ceramic Blank: Powder, Fabrication, and Properties . . . . . . . . 40 IV.1.1. Introduction 40 IV.1.2. Powder 45 IV.1.3. Cold Isostatic Pressing (CIP) 50 IV.1.4. Trimming of the Green Body 51 IV.1.5. Properties of the Green Body 52 IV.1.6. Heat Treatment 57 IV.1.7. Properties of the Blanks 62 IV.1.8. Colouration 82 IV.1.9. Characterisation of the Blank Fabrication Process 96 IV.1.10. Summary 101 IV.2. IV.2.1. IV.2.2. IV.2.3. IV.2.4.
Characterisation of Zirconia - Veneer Bilayer Structures . . . . Introduction Material and Methods Results and Discussion Summary
IV.3.
Fabrication of Multiple-Unit Dental Restorations by the Direct Ceramic Machining (DCM) Process . . . . . . . . . . Introduction Digitizing: Capturing the Shape Enlargement of the Digitized Shape: Compensation for Final Sintering Shrinkage Machining of the Ceramic Blank Sintering: Achieving Net-Dimensions and Materials Properties Veneering: Coating the Framework Summary
141 146 158 184 187
Preclinical Performance of the Bridges: Load Bearing Capacity and Reliability . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Introduction Material and Methods Results Discussion and Summary
191 191 193 199 211
IV.5.1. IV.5.2. IV.5.3. IV.5.4.
Clinical Performance of DCM fabricated Dental Restorations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Introduction Material and Method Results after the first year Discussion and Summary
216 216 217 222 224
IV.6.
References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 233
IV.3.1. IV.3.2. IV.3.3. IV.3.4. IV.3.5. IV.3.6. IV.3.7. IV.4. IV.4.1. IV.4.2. IV.4.3. IV.4.4. IV.5.
103 103 105 114 122 124 124 125
III
V.
Outlook . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 247 V.1.
VI.
Table of Contents
References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 249
Appendix . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 251 VI.1.
Chemical analysis of TZP powder lots . . . . . . . . . . . . . . . . . . . . . 251
VI.2.
Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 252
VI.3.
Calculation and Estimation of Errors . . . . . . . . . . . . . . . . . . . . . 254
VI.4.
TZ-3Y microstructure . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 258
VI.5.
Transmittance by Bouguer-Lambert Law . . . . . . . . . . . . . . . . . 259
VI.6.
Fabrication of plastic models for frameworks . . . . . . . . . . . . . . 259
VI.7.
Stress in bilayered structures . . . . . . . . . . . . . . . . . . . . . . . . . . . . 262
VI.8.
Anatomical tooth sizes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 264
VI.9.
Prototype Machine . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 264
VI.10. Series Machine . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 267 VI.11. References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 268
VII.
List of Figures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 271
VIII. List of Tables . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 277 IX.
Abbreviations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 279
X.
Curriculum Vitae . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 281
XI.
Work . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 283
Table of Contents
IV
1
I.
Summary
The Direct Ceramic Machining System (DCM)
I.1
Summary
Tetragonal zirconia polycrystals stabilized with 3 mol-% yttria (TZP) exhibits best mechanical properties, favourable aesthetic appearance and translucency. To date it is therefore one of the most promising candidates for dental application. Nevertheless this ceramic has not been used to a knowledgeable extent so far in prosthetic dentistry due to its difficulties in shaping. This drawback hinders the wider spread of its application although a strong market pull from patients and dentists already exists. Hence, the goal of this work was to develop and to make available a new system for the fabrication of dental restorations which solves the shaping difficulties. This new Direct Ceramic Machining (DCM) approach for TZP mills ceramic blanks in a porous, presintered and yet easily machinable state which are afterwards sintered to full density maintaining the required accuracy. The sintering shrinkage is compensated by enlarging the shape prior to machining. So far this approach has not been used in dental applications. Blanks were prefabricated and parameter sets were established. Easy machinability as well as high strength were achieved by adjusting presintering temperature as main trigger and compaction pressure as trigger of secondary importance. Homogeneity of the blanks on macro scale as determined from density distribution and on micro scale as determined by pore size distributions proved to be sufficient to guarantee the dimensional accuracy which is crucial for the dental application. Absolute dimensional accuracy was controllable within 19 µm (± 12 µm) after milling and sintering of a calibration speci-
Summary
2
men. A tolerance smaller than 0.09 % of length could be achieved which is more than sufficient for dental application. Coloration by infiltrating aqueous solutions of salts was proved to be feasible. However, colaration results need more detailed foundation, addressing especially the evolving colour, the thorough coloration of the complete blank and the effects on the properties of the frameworks. An experimental machine setup, tools, methods and appropriate parameter sets enable the reliable fabrication of dental bridges and crowns. The DCM procedure includes the digitizing of framework models using a mechanical probe mounted on a standard numerical controlled milling machine. The surface data are enlarged by 25 % to compensate for sintering shrinkage. Prefabricated blanks are machined using a modified standard milling machining with hard metal mills. Pre-clinical and clinical studies in their combination establish the feasibility of the DCM approach. Preclinical bend testing of three-unit TZP frameworks showed a more than two times higher load bearing capacity and two times higher reliability than other dental ceramic materials. A controlled clinical study in cooperation with University Zurich include 183 units of crowns, bridges with up to five units with two or more retainers in the anterior and in the posterior region in the lower and upper jaw. Inserted restorations cover the usual clinical indication bandwidth assuming that bridges with more than five units are rarely applied in clinical everyday life. Clinical study showed a hundred per cent success rate for more than two years of in vivo application. Hence it is concluded that this work proved the superiority of DCM fabricated dental restoration using a TZP framework. The results of this work are very promising and may enable for a commercialization of the DCM system. However, further adjustment and optimization will be needed.
3
I.2
Zusammenfassung
Zusammenfassung
Vollkeramische Systeme benutzen derzeit überwiegend Porzellan, Glaskeramik oder glasinfiltriertes Aluminiumoxid. Hochfeste, hochharte und hochzuverlässige Keramik hat nur begrenzte kommerzielle Verbreitung in der Zahnmedizin wegen der Schwierigkeiten ihrer Formgebung. Tetragonales polykristallines Zirkonoxid stabilisiert mit 3 mol-% Yttriumoxid (TZP) hat hervorragende mechanische Eigenschaften, ästhetisches Aussehen und Transluzenz. TZP ist der vielversprechendsten Kandidat für zahnmedizinische Anwendungen. Die prothetische Zahnheilkunde verwendet bereits TZP. Jedoch erfordert seine Hartbearbeitung im dichtgesinterten Zustand kostspielige Maschinen und Werkzeuge und benötigt lange Bearbeitungszeiten. Diese Nachteile verhindern die weitere Verbreitung, obgleich der Marktdruck von den Patienten und den Zahnärzten bereits existiert. Das Ziel dieser Arbeit war, ein System für die Herstellung von Zahnkronen und brücken zu entwickeln, zu evaluieren und verfügbar zu machen. Der Direct Ceramic Machining (DCM) Ansatz bearbeitet keramische Rohlinge in einem porösem, angesinterten und doch leicht bearbeitbaren Zustand. Danach wird unter Beibehaltung der zahnmedizinisch geforderten Genauigkeit dichtgesintert. Die Sinterschrumpfung wird kompensiert, indem die Form vergrößert aus dem Rohling herausgearbeitet wird. Dieser Ansatz wurde bisher in der Zahnmedizin nicht verwendet. Als Werkstoff wurde TZP ausgewählt. Innerhalb dieser Arbeit wurden die keramischen Rohlinge hergestellt, charakterisiert und Herstellparameter evaluiert. Gute Bearbeitungsfähigkeit und hohe Festigkeit wurden gleichzeitig erzielt durch die geeignete Festlegung von Ansintertemperatur und des Pressdruckes. Die Homogenität der Rohlinge wurde im Makrobereich als Dichteverteilung an Ringen gemessen und im Mikrobereich anhand der Porengröße und ihrer Verteilung bestimmt. Die Homogenität der Rohlinge ist ausschlaggebend für das massgenaue Sintern. Die absolute Maßgenauigkeit, gemessen an einem Kalibrierkörper, war kleiner als 19 µm (± 12 µm) nach Fräsbearbeitung und anschliessender Sinterung. Auf
Zusammenfassung
4
die Ausgangslänge bezogen wurde 0,09 % relative Toleranz erreicht. Für zahnmedizinische Anwendungen ist die Genauigkeit des DCM-Verfahrens ausreichend. Die Färbung durch Infiltration wässriger Kationlösungen wurde demonstriert, jedoch bedürfen die erreichten Ergebnisse hinsichtlich der entstehenden Farbe, der Durchfärbung der Rohlinge und der Auswirkungen auf die Eigenschaften der Grundgerüste weiterer Untersuchungen. Eine Experimentalmaschine, Werkzeuge, Hilfsmittel, Methoden und passenden Parameter erlauben die zuverlässige Herstellung der Zahnkronen und -brücken. Das DCM-Verfahren umfasst das Digitalisieren des Grundgerüstes mit einem Berührungssensor auf einer Standardfräsmaschine, die 25 %-ige Vergrösserung der Oberflächendaten zur Kompensation der Sinterschrumpfung und die Fräsbearbeitung der Rohlinge mit Hartmetall-Fräswerkzeugen. Die Machbarkeit des DCM-Ansatzes wurde durch erfolgreiche vorklinische und klinische Studien demonstriert. Vorklinische Biegebruchtests an dreigliedrigen TZP Grundgerüsten zeigten eine mehr als zweimal höhere Belastungsfähigkeit bei zweifach höherer Zuverlässigkeit im Vergleich zu anderen erhältlichen Dentalkeramiken. Eine kontrollierte klinische Studie in Zusammenarbeit mit der Universität Zürich umfasste 183 Zahneinheiten, die in Form von Kronen und Brücken mit bis fünf Gliedern, mit zwei oder mehr Verankerungen, im Front- und Backenzahnbereich im Unter- und Oberkiefer bei Patienten eingesetzt wurden. Der eingegliederte Zahnersatz umfasst die übliche klinische Indikationsbandbreite, unter der Annahme, dass Zahnbrücken mit mehr als fünf Gliedern selten in der täglichen Praxis anzutreffen sind. Während der mehr als zweijährigen Laufzeit zeigte die klinische Studie zeigt eine hundertprozentige Erfolgsrate. Innerhalb der vorliegenden Arbeit konnte die Überlegenheit des mittels dem DCMVerfahren hergestellten Zahnersatzes mit einem TZP Grundgerüst gezeigt werden. Die Ergebnisse dieser Arbeit sind erfolgversprechend für eine Kommerzialisierbarkeit des DCM-Systems, die allerdings noch weitere, über diese Arbeit hinausgehende Anpassungen und Optimierungen erfordert.
5
II.
Synopsis
Synopsis Abstract
All-ceramic dental restorations are available on the market and being used since the eighties. However, current dental ceramic materials have limited indication for allceramic restoration. Tetragonal zirconia polycrystals stabilized with 3 mol-% yttria in its dense state may be an alternative, as it exhibits more than two times higher strength and toughness. Despite of that, zirconia has not been widely used for all-ceramic restorations due to difficulties in shaping by traditional techniques, like casting and sintering, hotpressing or grinding in dense state. A new process called Direct Ceramic Machining (DCM) helps overcome these limitations by providing easy and accurate shaping of allceramic dental restorations. This chapter introduces the topic, the goal and the approach of this work.
Synopsis
II.1
6
Introduction
Ceramic materials for dental restorations were first invented in the 18 th century by the Parisian pharmacist Alexis Duchâteau (1714 - 1792) [1]. He considered porcelain as an alternative to traditional ivory-made restorations in order to avoid the drug-caused discoloration typical of ivory. In collaboration with the Parisian dentist Nicolas Dubois de Chémant he developed the first known application of porcelain in restorative dentistry. De Chémant conducted further research into this invention, earned his first PhD in 1798 and was awarded patents. Based on this work, Giuseppangelo Fonzi (1768 - 1840) invented single porcelain teeth joint together with gold supporting structure. Until 1960 most advances in application of porcelain were related to improvements in aesthetics particularly in translucency - and to fabrication methods of teeth in dental practices [2]. In 1962, Weinstein, Katz and Weinstein [3], and Weinstein and Weinstein [4] were awarded their US patents for a gold alloy and a feldspathic porcelain for the fabrication of porcelain-fused-to-metal (PFM) restorations, respectively. Since then the use of porcelains has been broadened in restorative and prosthetic dentistry [2]. Today, dental bridges are mostly metal-porcelain composite structures and consist of a metallic framework for load bearing which is coated with porcelain for aesthetic appearance. Porcelain-fused-to-metal (PFM) restorations have set the standard for multiple teeth restoration in the past decades. They account for approximately 80 % of all fixed restorations and show a clinical survival rate of 95 % after 7.5 years or more [5]. However, despite this success the use of metal in the oral cavity has come under dispute in recent years due to its biological incompatibility risk. In addition, PFM shows severe limitations concerning aesthetic appearance due to the metallic framework and does not satisfy market needs for enhanced aesthetics. Therefore, all-ceramic restorations are of great importance and would be of high clinical value.
7
Synopsis
All-ceramic restorations are available on the market. Most of them use feldspathic porcelain, leucite reinforced porcelain, glass-infiltrated porous alumina, and glass-ceramics [5, 6]. The shortcomings of these materials are low bend strength and low toughness (fig. 1), which lead to design restrictions, non-reliability and complicated multi step manufacturing procedures. Due to this, use of dental ceramics has been limited to crowns, onlays and inlays. During the past decade special attention has been paid to improvement of the mechanical properties of ceramics, and to shaping technologies. The intention behind this was not only to extend the indication bandwidth of all-ceramic restorations, but also to achieve sufficiently low failure rates [2, 5].
Toughness [MPa m1/2]
10
High- Tech Ceramics
8
Zirconia (present work)
Glass-Infiltrated Alumina
6
In-Ceram Zirconia
Glass-Ceramic
In-Ceram
4 Conventional Porcelain
Empress2
(Vita-Celay) Alumina
Dicor MGC
2
In-Ceram
figures/materials.wmf
Empress MK II Omega
0
0
200
400
600
800
Bend Strength [MPa] Fig. 1:
Dental ceramic materials properties used for restorations [7-13].
1000
Synopsis
8
The intra-oral environment requires ceramic structures to withstand cyclic loads and water [2, 14]. Wide temperature (5 °C to 55 °C) and pH (0.5 to 8.0) shifts are another intra-oral reality, however probably represent a secondary concern for ceramics. Average chewing forces during normal mastication depend on the region of the mouth and are reported in a wide range from 40 N to 440 N. Higher forces can readily be reached (500 N to 880 N) but are more likely associated with parafunctional habits such as clenching and grinding (bruxism) than with normal chewing. Actual tooth-to-tooth contact is more likely during parafunctional habits than during normal chewing of food, and many believe that the food bolus contributes greatly to the wear of teeth and restorations. Individual areas of contact between opposing teeth are in the range of 1 mm2 to 4 mm2 placing local contact pressures in the general range of 2 MPa to 150 MPa. The number of chewing cycles per day is reported to be around 800 to 1400. Water is obviously available to all external ceramic surfaces exposed to saliva. However, all internal ceramic surfaces, as well, are exposed to water because dentin (inner tooth structure), micro leakage and diffusion through dental cements are likely sources of water access. Nevertheless, the oral hostility offers a challenge to ceramic materials which may not be simulated in-vitro but approached by clinical studies. New ceramics and ceramic composites potentially meet these market needs for dental bridges. One of the most promising candidates is tetragonal zirconia polycrystals (TZP) stabilized with 3 mol-% yttria (3Y-TZP) [15]. In its dense sintered state it exhibits much higher strength and toughness than all other dental ceramics. Moreover, TZP is interesting for dental applications as it fulfils high aesthetic demands in colour and translucency. Furthermore its biocompatibility is already established from its in-body use for femoral head balls in hip replacements [16, 17]. Despite its favourable properties TZP has only been used in very limited quantities for all-ceramic dental bridges due to its difficulties in shaping by usual techniques like casting and sintering, hot pressing or grinding in its dense form. Hence, the fabrication of all-ceramic restorations with a zirconia framework needs a sophisticated fabrication system which combines materials, devices, tools and methods. For this work, we chose TZP as framework material veneered with a feldspathic porcelain adjusted to TZP. The machining uses shaping tools for the TZP framework which are appropriate for the frameworks shape, the material and its properties. The me-
9
Synopsis
tods how to shape the material with devices and tools is essential. Requirements for such a system combine the ease of handling by dental technicians, suitability for crowns and multi-unit bridges, the possibility to fabricate individual and complex shapes with high accuracy and reliability, as well as fast and economic fabrication. Integration of the system in the patient treatment process requires close and smooth cooperation between patient, dentist and dental technician.
II.2
Goal and Approach
The goal of this work was to develop and make available a new system for the fabrication of dental restoration using a zirconia framework. The system combines materials, operating devices, tools and methods. Furthermore the system had to be evaluated in pre-clinical analysis in comparison to other commercially available systems. The clinical feasibility of the system had to be demonstrated by in-vivo studies. However, these in-vivo studies required positioning the system in the clinical treatment procedure, and therefore included the definition of interfaces to dental technicians and dentists. The Direct Ceramic Machining (DCM) approach is shown in fig. 2. The main idea is to simplify the shaping of ceramic frameworks by machining a soft pre-sintered blank, and to sinter it to its full density afterwards. This process requires the sintering shrinkage to occur homogeneously in all spatial directions in order to guarantee dimensional accuracy and to avoid final hard machining in the dense sintered state.
Synopsis
10
figures/approach.jpg
Fig. 2:
Direct Ceramic Machining (DCM) system approach [18].
A system capable to operate the DCM process needs a digitizing device, a computer for points capturing, enlarging and calculating the tool paths, a machining device, a sintering furnace, tools for digitizing and machining, homogeneous blanks, and veneering porcelain with appropriate devices for applying the veneer. The outer shape of a framework model for dental bridges is mechanically digitized using a 3D sensor. Digitizing produces a digital representation of the surfaces consisting of many data points. These surface data points are isotropically enlarged in all spatial directions for compensating of the final sintering shrinkage. Tool path information is generated from the enlarged surface data set and transferred to a milling machine. The framework is milled from a homogeneous porous ceramic blank of TZP, the properties of which are adjusted to be easily machinable. The result is an enlarged yet porous framework. During final sintering to full density the enlarged framework shrinks to the original dimensions maintaining the required accuracy with no further hard machining needed. In the final step, the framework is coated with veneer porcelain to meet the patient’s aesthetical requirements in colour and translucency.
11
Synopsis
The approach has to meet two key requirements, which are high accuracy and the ability to machine thin-walled gracile frameworks with sharp cervical edges out of mechanically weak blanks. In addition, tool wear needs to be low for cost and accuracy reasons. Accuracy needs to be smaller than ± 25 µm on 30 mm, which is the average length of a three-unit framework. The achievable accuracy is mainly determined by the homogeneity of sintering shrinkage, which is a function of the density homogeneity of the blanks. Hence, homogeneous blanks are a prerequisite for the system to work successfully. The highest challenge for the process lies in machining of the slender edges at the cervical margin, which is largely influenced by the properties of the blanks, in particular easy machinability and material strength in order to withstand the cutting forces.
II.3
Outline of the Thesis
This thesis was motivated by the fact that there is no solution for shaping zirconia dental frameworks available which would fulfil the requirements necessary for broad use of all-ceramic dental restoration. The second chapter will provide evidence for this. It summarizes the state of the
art of fabrication systems for dental bridges, and discusses their advantages, disadvantages and limitations. Chapter three outlines the DCM system approach and therefore occupies the central position in this work. Its subchapters are dedicated to the blanks, the framework and veneer material evaluation, the DCM steps to dental restorations, and its pre-clinical and clinical evaluation. We start with the fabrication and characterisation of the blanks because it is the crucial prerequisites for establishing the DCM process. Then, the veneer porcelain especially for zirconia frameworks was developed and mechanically analysed using zirconia porcelain bilayered structures. At that point of the work the materials are analysed and blanks may be fabricated. In the third subchapter the tools, methods and parameters for the fabrication of zirconia frameworks and application of the veneer por-
Synopsis
12
celain are established. Pre-clinical simulation and experimental tests using pure and veneered three-unit frameworks prove DCM’s principal feasibility. The clinical evaluation of the system is the final step to prove its clinical feasibility. An outlook and a summary conclude this work.
II.4 [1]
References
Ring, M.E.: "Geschichte der Zahnmedizin.", 1997, Köln: Könemann Verlagsgesellschaft mbH.
[2]
Kelly, J.R.: "Ceramics in Restorative and Prosthetic Dentistry." Annual Reviews
Materials Science, 1997. 27: p. 443-468. [3]
Weinstein, M., Katz, S., and Weinstein, A.B.: "Fused Porcelain-to-Metal Teeth.", 1962, United States Patent Office, US 3,052,982.
[4]
Weinstein, M. and Weinstein, A.B.: "Porcelain covered Metal-Reinforced Teeth.", 1962, United States Patent, US 3,052,983.
[5]
Anusavice, K.J.: "Development and Testing of Ceramics for Dental Applications."
Ceramic Transactions, 1995. 48: p. 101-124. [6]
George, L.A. and Eichmiller, F.C.: "Dental Applications of Ceramics." in Biocer-
amics: Materials and Applications., G.S. Fischman, A. Clare, and L. Hench, Editors. 1995, The American Ceramic Society: Indianapolis IN. p. 157-172. [7]
Rieger, W.: "Medical Applications of Ceramics." in High-Tech Ceramics: View-
points and Perspectives., G. Kostorz, Editor. 1989, Academic Press: London. p. 191-228. [8]
Geis-Gerstorfer, J., Kanjantra, P., Pröbster, L., and Weber, H.: "Untersuchung der Bruchzähigkeit und des Risswachstums zweier vollkeramischer Kronen- und Brückensysteme." Deutsche Zahnärztliche Zeitung, 1993. 48: p. 685-691.
[9]
Seghi, R.R., Denry, I.L., and Rosenstiel, S.F.: "Relative Fracture Toughness and Hardness of New Dental Ceramics." Journal of Prosthetic Dentistry, 1995. 74(2): p. 145-150.
[10] Seghi, R.R. and Sorensen, J.A.: "Relative Flexural Strength of Six New Ceramic Materials." International Journal of Prosthodontics, 1995. 8: p. 239-246.
13
Synopsis
[11] Lüthy, H.: "Strength and Toughness of Dental Ceramics." in CAD/CIM in Aesteth-
ic Dentistry, W.H. Mörmann, Editor. 1996, Quintessence: Chicago. p. 229-239. [12] Lüthy, H.: "Personal Communication.". 2000: Zurich. [13] Oh, S.-C., Dong, J.-K., Lüthy, H., and Schärer, P.: "Strength and Microstructure of IPS Empress2 Glass-Ceramic after Different Treatments." International Journal of
Prosthodontics, 2000. 13(6): p. 468-472. [14] Kelly, J.R.: "Clinical Failure of Dental Ceramic Structures: Insights from Combined Fractography, In Vitro Testing and Finite Element Analysis." Ceramic
Transactions, 1995. 48: p. 125-137. [15] Gauckler, L.J.: "Processing and Properties of Advanced Structural Ceramics." in
High-tech ceramics: viewpoints and perspectives., G. Kostorz, Editor. 1989, Academic Press: London. p. 59-105. [16] Richter, H.G., Burger, W., and Osthues, F.: "Zirconia for medical implants - the role of strength properties." in International Symposium on Ceramics in Medicine. 1994. Turku, Finland: Butterworth-Heinemann Ltd. [17] Drouin, J.M. and Cales, B.: "Yttria-Stabilized Zirconia Ceramic for Improved Hip Joint Head." in International Symposium on Ceramics in Medicine. 1994. Turku, Finland: Butterworth-Heinemann Ltd. [18] Filser, F., Lüthy, H., Schärer, P., and Gauckler, L.J.: "All-Ceramic Dental Bridges by Direct Ceramic Machining (DCM)." in Materials in Medicine, M.O. Speidel and P.J. Uggowitzer, Editors. 1998, vdf Hochschulverlag AG an der ETH Zürich: Zürich. p. 165-189.
Synopsis
14
15
III.
State of the Art
State of the Art: Fabrication of AllCeramic Restorations Abstract
A variety of systems are applied for the fabrication of all-ceramic dental restoration. They include systems using basic ceramic shaping methods such as hot-pressing, slip casting, compaction of powder, machining in hard and soft state of the material, and some other approaches. This chapter which describes the systems used in the dental application to fabricate all-ceramic tooth restoration is based on patent research as well as on a market survey. In each subchapter the description of the principles precedes the analysis of the systems. Known systems exhibit limitations concerning fabrication of the type of restoration, the material, the reliability, the simplicity or the cost.
State of the Art
III.1
16
Introduction
Dental restorations made of ceramic material are in demand since many years. Ceramic inlays, onlays and crowns were successfully introduced in the dental market. Bridges especially for molar-region use require high-strength, high-toughness ceramic material. Shaping ceramics to the complex three-dimensional freeform shape of the restoration is very challenging. Furthermore each restoration is one-time individually fabricated and used but once for a patient. This requires an accurate, fast, cheap, reliable and simple fabrication procedure for high-strength ceramics. For shaping ceramics a variety of technologies have been investigated. According to DIN 8580 [1] two basic principles distinguish the shaping by material removal from a stock and the shaping by material addition. Turning, milling, and grinding processes are in the first group whereas the second group comprises the classical ceramic forming methods like slip casting or pressing as well as the more recent rapid prototyping techniques. König et al. [2] and Klocke et al. [3] presented good overviews on material removal technologies possible for ceramic materials (but without regard on dental application): machining of green and presintered ceramics, traditional hard machining processes like grinding, lapping, polishing, honing, ultrasonic machining, electrodischarge machining, laser beam machining, and laser assisted machining. Richerson [4] and With [5] described the forming methods that are well known for ceramic materials where parts are made by consolidation of powder, dough, or slips by compacting (pressing), slip casting, hot pressing, or injection moulding or by casting techniques. Rapid prototyping present recent approaches dealing with fast production of a small number of complexshaped ceramic parts [6, 7]. Especially for fabricating ceramic parts three dimensional printing, selective laser sintering and freeform powder moulding represent interesting technologies. CAD/CAM devices were widely addressed in literature [8-14]. These were analysed concerning the types of restorations and the material fabricable (indication width), the steps supported along the restoration fabrication (process width), their place of installation in dental practice or dental technician’s laboratory, central or de-central fabrication facilities, digitizing abilities, machining abilities, the designing capabilities, in-
17
State of the Art
vestment costs, and the properties of the fabricated restorations. Despite the fact that these properties are absolutely important for decision makers they will change rapidly with time. Hence the CAD/CAM devices as available on the market will be assigned to basic ceramics shaping technologies to describe and analyse their potentials. The goal of this chapter is to give an overview on shaping systems, however the overview it will be focused towards systems enabling the fabrication of ceramic dental restorations. They are structured along the basic ceramic shaping technologies such as hot and cold pressing, slip casting, hard and soft machining and some other approaches. Within each subchapter the shaping principle is described followed by the systems based on that technology.
III.2
Hot Pressing - Systems
Hot pressing is a simultaneous forming and sintering procedure under overpressure [4, 5, 15]. In general uniaxially and isostatic pressure application is distinguished. Hot pressing starts with the fabrication of the moulds by e.g the lost-wax technology. In the moulds the powder material or a prefabricated ingot is placed. When mould and ingot are heated to sintering temperature the pressure is applied via a ram and the material fills the mould cavity. The mould refractory material is the most important material in the hot press. It must withstand the temperature, the thermal stresses, the stresses by pressure application at elevated temperature with minimal deformation and being chemically inert to framework material. The mould must be fabricated easily and accurately for each individual framework. Hot pressing of complex shapes like frameworks is difficult. Non-uniform cross-sections, non-uniform filling, may lead to flaws or pores.
State of the Art
18
Hot pressing is used in dental application since years and many systems use this method. Typical representatives are Ivoclar’s IPS Empress®1 and IPS Empress2®1 systems, Dentsply Ceramco’s Finesse All-Ceramic®2 system, and Degussa Ney Dental’s Cergogold®3 system. According to the manufacturer the raw material varies in its material composition, its processing parameters such as temperature, time and pressure, and the crystalline phases evolving during sintering of the restoration. Typical composition of IPS Empress include 17 to 23 wt-% of Al2O3, 10 to 14 wt% of K2O, 3 to 7 wt-% of Na2O, 58 to 63 wt-% of SiO2 as main components, additives such as B2O3, BaO, CaO, CeO2, TiO2 and 0.5 to 2 wt-% of pigments [16-18]. The ingots are pre-cerammed by the manufacturer. Leucite crystal (K2O • Al2O3 • 4 SiO2) content is in the range of 40 to 60 vol-%. Depending on its composition the glass transformation temperature is in the range of 625 to 650 °C. Mould and ingot are warmed-up slowly to 850 °C, and tempered for 90 min to get a homogeneous temperature distribution. After heating to 1100 °C a slight pressure is applied onto the piston until its movement per minute is smaller than a preset treshold value. Afterwards the mould is destroyed, the restorations are glasspearl-blasted, and the sprues are removed. Finesse4 and Cergogold5 are in the same phase system and therefore are similar to the IPS Empress system. IPS Empress2 consists of up to 5 wt-% Al2O3, up to 13 wt-% K2O, up to 6 wt-% La2O3, up to 13 wt-% Li2O, up to 5 wt-% MgO, 3 to 7 wt-% Na2O, up to 11 wt-% P2O5, 57 to 80 wt-% SiO2, up to 8 wt-% ZnO as main components, up to 8 wt-% other additives and pigments. Lithium disilicate Li2Si2O5 crystal content is more than 60 vol-% with additions of a secondary lithium orthophosphate Li3PO4 crystals [17, 20, 21]. Its glass transformation temperature is 535 °C. Pressing temperature is decreased to 920 °C. 1. 2. 3. 4.
5.
Ivoclar AG, Bendererstrasse 2, FL-9494 Schaan, Principality of Liechtenstein DENTSPLY Ceramco, Six Terri Lane, Burlington, NJ 08016, USA Degussa-Ney Dental, Inc., 65 Dudley Town Road, Bloomfield, CT 06002, USA Composition is 6 to 12 wt-% Al2O3, 5.5 to 10.5 wt-% K2O, 6 to 12 wt-% Na2O, 40 to 65 wt-% SiO2 as main components, additives such as B2O3, CaO, CeO2, Li2O, and pigments according to [19] and actual composition according to Dentsply’s marketing division is 10 to 15 wt-% Al2O3, 5 to 10 wt-% K2O, 5 to 10 wt-% Na2O, 60 to 65 wt-% SiO2, and additives of 1 to 3 wt% B2O3, 0 to 2 wt-% CaO , 0 to 1 wt-% BaO , 0 to 1 wt-% CeO2 , 0 to 1 wt-% TiO2. Composition according to: 12 to 25 wt-% Al2O3, 0.1 to 2.5 wt-% CaO, 7 to 18 wt-% K2O, 0.5 to 25 wt-% Na2O, 50 to 80 wt-% SiO2 as main components, additives such as BaO, CeO2, La2O3, Li2O, MgO, Sb2O3, SnO2, SrO, TiO2, ZnO, ZrO2, and pigments.
19
State of the Art
Research conducted at the Swiss Federal Institute of Technology for hot pressing of Y-TZP or Y-TZP with additives of various glasses produced unsatisfactory results either due to the microstructure and mechanical properties of the sintered body [22-24] or due to unsatisfactory moulding materials [25-27]. Advantages of hot pressing include the reduced cycle time compared to pressureless sintering, the reduced sintering temperature and therefore reduced grain growth achieving higher strength, the minimized residual porosity, and that less additives are required. Most of the hot pressing systems achieve a good accuracy and the framework fits well to the prepared teeth. The disadvantages include expensive equipment, lost moulds including the expense of mostly manually performed moulds fabrication, and the complicated error-prone process. The used materials are low fusing and include a high percentage of amorphous phase which is responsible for their high translucency. However, they exhibit low strength and toughness because of the size and distribution of flaws created during processing.
III.3
Casting - Systems
Casting systems fabricate the dental restorations by casting a fluid melted glass in a mould, cooling it down and afterwards ceramming this object to get the desired properties. The requirements for the refractory material are in principle equivalent to hot pressing systems. The mould is fabricated using the lost-wax technology. Representative for this technology are Dentsply Int.’s and Corning’s Dicor® [13, 28]. Dicor consists of K2O, MgF2, MgO and SiO2 as main components with additions of Al2O3 and ZrO2. It is a micaceous glass-ceramic with 55 % crystalline phase of tetrasilicic fluoromica K2Mg5Si8O20F4. The amorphous glass ingot is melted at 1350 °C, centrifugal cast, and cooled down. Afterwards the cast glassy objects were cerammed at 1075 °C for 6 hours.
State of the Art
20
Casting processes are well known to the dental technician - even if they mostly deal with metals. Disadvantages of those casting systems include the expensive equipment of a special centrifugal casting device, the design and fabrication of lost moulds which require special skills, and the time consummation of heating the moulds and of the ceramming. The ceramming step is crucial because it controls the amount of crystalline phase in the restoration, and therefore determines its mechanical properties. Despite all disadvantages, the cerammed restoration exhibits an outstanding translucency.
III.4
Slip Casting - Systems
Slip casting involves ceramic particles suspended in aqueous or organic solvents which are cast into porous moulds [4, 5, 29]. Moulds are usually made of plaster of Paris and remove the solvent by controlled capillarity or additional external pressure. Representatives for this technique is Vita’s In-Ceram® and its automated version called Wolceram. In-Ceram uses fine-grained alumina powder with grain sizes from 2 to 5 µm [3032]. From the silicon impression the dental technician fabricates a plaster master model representing the positive situation in the patient’s mouth. The special plaster possesses an adjusted expansion coefficient. The ceramic slip is manually dabbed with a brush on the master model to form a framework and contains alumina powder and a special fluid which is homogenized using an ultrasonic horn. Where the slip touches the plaster wall the fluid is removed through capillary effects. Work-over using a scalpel is possible for optimization of the framework form. Then the dental framework on the plaster model is partially sintered at 1120 °C for 2 hours. The framework shows only minimal shrinkage whereas the plaster abutment teeth shrinks more. The porous alumina framework can thus be easily removed from the plaster model and is strengthened afterwards by glassinfiltration. Different colours require different infiltration glasses. Glass powder is mixed with distilled water and applied on the framework and fired for 4 hours at 1100 °C. Excess glass has to be removed by sandblasting. In the last step the veneer porcelain is applied.
21
State of the Art
The Wolceram process [33] uses an automated slip casting and milling process on the basis of the digitized data. A plaster model of the abutment tooth is fabricated as described before. Then, the slip is applied on the plaster model of the abutment tooth to shape the cavity, it is dried and afterwards milled to shape the outer surface. The firing and infiltration procedure is as described for In-Ceram. Complicated shapes can be fabricated using slip casting at relatively low cost. But the slip casting is a complex process requiring careful control of each single step: viscosity, settling rate, air bubbles, casting rate and many more. Therefore, flaw-free fabrication of restoration with a brush in the dental laboratory is a difficult, non-reliable task. An automated version of the slip casting technology is available, but it does not omit the disadvantages of the slip casting process itself. Moreover slip casting is a slow process. To handle slips is more difficult than handling solid preforms. The attempt to transfer the In-Ceram technology to Y-TZP leads to weaker mechanical properties than for alumina [34].
III.5
Compaction of Powder
Compaction of powder is a widespread technology for fabricating simple shaped ceramic parts. Powder is filled in a mould and compacted to a green body by applying pressure [4, 5, 15]. Different types of compaction technology are known depending on the direction of compaction (uniaxial or isostatic), the state of the powder (wet or dry), and the temperature (cold or hot pressing). Subsequent sintering of the cold compacted body leads to shrinkage dependent on green density. Enlarging the moulds to compensate for final firing shrinkage enables near-net shape for parts. Nobel Biocare’s 1 Procera ® successfully uses this technology for the fabrication of dental crowns and bridges of alumina.
1.
Nobel Biocare AB, Box 5190, 402 26 Göteborg, Sweden
State of the Art
22
Procera uses the digitizing data for an abutment to produce enlarged moulds [35, 36]. The alumina powder is filled in the mould cavity and high-pressure compacted. The compacted bodies were adjusted along the cervical margin and then sintered to full density at 1550 °C for one hour. Three-unit bridges are fabricated of three single units which are joined together by glass soldering [37]. Afterwards the veneer porcelain is applied. Nobel Biocare Procera technology uses a centralized manufacturing facility in combination with decentralized digitizing devices. Dental laboratories do the digitizing of the plaster abutments, the computer-aided design and then transmit the data to Nobel Biocare’s facilities for fabrication of the framework [38]. The Procera system has successfully proved its ability to reliably fabricate crowns [39]. Nobel Biocare has set up 980 scanners worldwide up so far and fabricated 650’000 single-unit restorations per year for more than 300’000 patients since 1995 [40]. The frameworks are fabricated using industrial equipment for series production. However, the achievable accuracy is said to be limited [41]. Possible shapes depend on the moulds. Bridges are not one monolithic part but three single joined-together units. Disadvantages include the expensive fabrication devices requiring centralized facilities which cause a decrease in value-adding steps and in flexibility at the dental laboratory.
III.6
Hard Machining - Systems
Hard machining systems use abrasives for micro removal of particles from a dense sintered stock [42, 43]. Hard machining is distinguished by the state of the abrasives. Honing, lapping and polishing use loose abrasive grains whereas grinding uses grains embedded in a matrix and applied to a support, e.g. wheel or a cylindrical rod. The latter one exhibits in general higher material removal rates than the first ones. In dental application bound abrasives are preferred because a restoration has to be machined out of pre-
23
State of the Art
fabricated stock material. Representatives in dental application are systems marketed by Advance1, DCS Dental2, Decim3, Girrbach Dental4, Sirona5, and Elephant’s Cicero® 6. Their commercially available systems are using grinding technology in hard state, but may differ in the material used for the framework, the digitizing and the machine. The Cadim® system by Advance uses porcelain and machinable ceramics in form of prefabricated blanks to grind. The system operates with a mechanical digitizing unit (Renishaw7). Design of the frameworks is done interactively. The system is designed for laboratory use and is not yet on the market. The President DCS® - system by DCS Dental and the Digident® - system by Girrbach Dental, respectively use dense sintered ceramics such as TZP or Vita’s In-Ceram® or Zirconia ® blanks to grind. Both systems possess a very stiff design to enable for grinding of TZP. Both systems include separate digitizing and machining devices that are linked by computer-aided design for designing the frameworks with which the tool path informations are generated and the data are transmitted to the machining device. The President DCS includes a laser scanner whereas the Digident uses line grid projection technique. The plaster model of the mouth is digitized. Both systems enable also a centralized business solution, where the scan data are transmitted to a milling facility and the frameworks are shipped to the dental laboratory. Decim® system by Decim company uses dense sintered ceramics such as Denzir®, a yttria stabilized zirconia (supplied by Norton Desmarquet8). Decim’s machining process (patent pending) uses high-speed under-water grinding [44]. The system consists of three separated devices for laser digitizing, designing and machining and is designed for use in dental laboratories acting as service centers.
1. 2. 3. 4. 5. 6. 7. 8.
Advance Co. Ltd., Dental Department Cadim Division, Nihonbashi Kobuna-cho, 7-7 Tokyo 103-8354, Japan DCS Dental AG, Gewerbestrasse 24, 4123 Allschwil, Switzerland Decim AB, Box 733, 931 27 Skellefteå, Sweden Girrbach Dental GmbH, Dürrenweg 40, 75177 Pforzheim, Germany Sirona Dental Systems GmbH, Fabrikstrasse 31, 64625 Bensheim, Germany Cicero Dental Systems BV, Elephant Dental B.V., Hoorn, Netherlands Renishaw PLC, Wotton-under-Edge, Gloucestershire, GL12 8JR, K Norton Desmarquet, 48 rue des Vignerons, 94685 Vincennes Cedex, France
State of the Art
24
The Sirona Cerec® - system by Sirona Dental uses dense machinable ceramics in form of prefabricated blanks such Cerec Bloxx®, Vita’s Vitablocs ®, Vita Mark II® , Dentsply’s Dicor MGC® or Ivoclar’s ProCad®. Furthermore Cerec enables to grind Vita’s In-Ceram® or Zirconia® blanks. The system in its recent version, Cerec III, consists of separate intraoral imaging device, CAD/CAM workstation and a combined device for laser digitizing and machining. Hence, the systems enables to acquire data directly from the tooth in the patients’s mouth or from a plaster model of the situation in the mouth. The machining device possess two equal machining spindles which are grinding the occlusal and the cavital side of the restoration at the same time. The machining device is to weak for hard machining of dense TZP. Cicero® - system fabricates the inner cavity of the prepared tooth out of a thermalexpansion-adjusted refractory [11]. On this a first layer of Synthoceram (66 wt-% Al2O3, 2 wt-% K2O, 1 wt-% Na2O, 27 wt-% SiO2, 4 wt-% ZrO2) is compacted, sintered and excess material is removed by grinding. In order to get the complete crown further layers are applied using the same procedure as for the first layer. Therefore Cicero-system enables the colour layering in order to mimic natural teeth’s natural appearance. The system consists of a digitizer for scanning the plaster model, a software for designing the restoration, and fabrication facilities. The idea is to get the silicon impressions and to ship the ready-to-use restorations back to the dentist. The Cicero process seems to be too complicated for easy use in dental laboratories, it requires expensive equipment and special knowledge. Traditional hard machining is very well known and draws its experience from decades of application and research. Hard machining produces very accurate parts possessing surfaces with low roughness if polished. But hard machining processes are slow, require stiff, well damped, high powered and accurate machines, special tools and cooling media [42]. These are some of the reasons why hard machining is very expensive. Newer approaches tend towards high speed grinding of ceramics. This technology allows higher cutting speed (150 m/s in relation to conventional grinding 25 m/s), feedrate, and cutting depth [45]. High speed grinding is said not to decrease strength of the machined ceramics as conventional grinding does [46] due to creation of subsurfacial flaws [43, 47-51]. Moreover, high-speed grinding is advantageous compared to conventional grinding in speed and, accordingly machining time. The goal of hard machining is to
25
State of the Art
fabricate the net-shape. Nevertheless, tooling for grinding is very limited towards small tool diameters due to high cutting forces and high temperatures. As a consequence grinding the material in hard state limits the ability to fabricate gracile and complex surfaces, especially cavities and depressions as required by dental application. Therefore achieving net-shape for the frameworks often requires some manual over-working by the dental technician.
III.7
Soft Machining
Porous high-tech ceramic materials are easy to machine with low cutting and feed forces compared to its dense sintered counterparts. Hence, ceramic blanks enable to fabricate restorations in combination with machining devices. Representatives of commercially available blanks are Vita’s In-Ceram Alumina® (Al2O3) and In-Ceram Zirconia® (Al2O3 with addition of ZrO2). Grinding of In-Ceram blanks has been already used for years by Mikrona Technologie’s Celay® manually operated copy grinding machine1. In fact the previously mentioned machines for hard machining such as the President DCS, the Cerec, and the Digident - system may also be of advantage for soft machining. Other processes like the Wolceram process make also use of this approach. Wolceram removes the dried excess slip to shape the occlusal surface of the restoration. The Celay machine requires to fabricate manually a positive plastic model of the framework. The blank and the model are clamped in the machine [52]. A 1:1 copy of the plastic model is created by sliding a sensing tool on the model’s surface which leads the grinding tool to cut the framework of the blank. Then the porous In-Ceram framework is glass-infiltrated to gain its final mechanical properties. Soft machining produces only low forces, but it uses the same grinding tools and technology as in case of hard machining. In any case small statistically aligned but geometrically non-defined micro cutting edges are responsible for material removal. Usually grains of hard material like diamond, corundum or cubic boron nitride are bound to a
1.
Mikrona Technologie AG, Spreitenbach, Switzerland
State of the Art
26
tool holder or embedded in a matrix and create the cutting edges. Smaller grain size lead to higher accuracy but material removal rate drops and vice versa. Due to the low forces soft machining tools have the advantage to allow even smaller tool sizes than used for hard machining.
Soft machining has been also understood as the last step of shape modification as this is true in case of hard machining. Furthermore it has been said to satisfy near-net
shaping requirements, and to enable for only 3 % relative dimensional tolerance for normal procedure if final firing shrinkage is compensated [41, 53]. Even precision processes may achieve 1 % relative dimensional tolerance. Lacking the ability for net-shape
fabrication and low dimensional tolerances are two major drawbacks to the dental application of soft machining. König et al. [2] as well as Klocke et al. [3] considered green machining as a special type of soft machining as flexible and cost-effective process for small-to-medium batch sizes. But, in their opinion so far no systematic study of green machining is known and the know-how is mainly based on empirical data. Major drawbacks were seen in the fragility of the green bodies which have the disadvantage in transmission of cutting forces and fragility (see also [4]). Additionally, adequate machines, tools and clamping devices are lacking for soft machining technology. Both references consider green machin-
ing as a means for near-net-shaping. Lacking the ability for net-shape fabrication and low dimensional tolerances are two major drawbacks. From this one must conclude that green machining is not adequate for dental application. The Vita-Celay system uses abrasive machining of soft alumina bodies followed by glass infiltration [52, 54]. The main goals for performing soft machining are to shorten the fabrication time and to avoid the costly finish machining in the hard state.
27
III.8
State of the Art
Other Approaches
Ultrasonic Machining Ultrasonic or sonoerosive machining is a lapping process supported by high-frequency longitudinal ultrasonic vibration and is applied in the hard state of the ceramics. It is said to permit a “(...) free-selection of geometrical shapes within a wide range of capabilities (...)” [3]. The technology is described in [55], its dental application is mentioned in [56]. Ultrasonic machining is suited for materials with brittle failure and stress intensity factors lower than 8 MPa√m, and therefore limits the materials selection. However, the ultrasonic machining is judged as very complicated due to the necessity to fabricate metallic tools and it is slow due to small material removal process. A broader dissemination in dentistry is not yet known.
Electric Discharge Machining (EDM) EDM removes small portions of material by successive electrical discharges between electrically conductive materials [3]. Electrical conductivity limits the ceramic materials to be machined. Shapes to be manufactured are determined by the geometry of the tool or by tool kinematics and, therefore complex shapes are possible using EDM technology. EDM is said to be fast and accurate. But, EDM seems to require elaborated machines and tools making it not the first choice for dental application. Nevertheless it was used for fabrication of titanium crowns, for example by Andersson [57].
Laser beam machining Laser beam machining removes little portions of material by vaporization using high energy pulsed lasers. Thus, thermoshock stability and low thermal conductivity are prerequisites for the material. This technology has the ability of creating complex shapes, however, it’s application for fabricating dental restoration is unknown.
State of the Art
28
Laser-assisted machining Laser-assisted machining improves machinability of hard materials by locally heating up locally to high temperatures at the place of cutting thus giving the material a partial plastic behaviour [3, 58]. Laser-assisted turning already showed feasibility if the stock material has a low thermal expansion, and high thermocycling and thermoshock stability. Machining at elevated temperatures requires special tools. Despite its potentials for simple contours no single application for fabricating dental restoration with their complex geometry is known up to now.
Shrinkage-free Ceramic Materials Shrinkage-free ceramic materials will minimize dimensional changes during final sintering [59, 60]. Hennige et al. [60] processed ZrSiO4 via reaction sintering of ZrO2 and SiO2 at temperatures higher than 1100 °C. Shrinkage during sintering was compensated by an volume increase at temperatures in the range of 650 °C to 1000 °C by decomposition and oxidation of ZrSi to the reaction sintering agents ZrO2 and SiO2. To get zero shrinkage green density has to be carefully controlled. Lee et al. [59] modified the In-Ceram system by addition of aluminium. During sintering oxidation of approximately 3 vol-% aluminium compensates for the shrinkage of the alumina matrix. Avoiding the shrinkage compensation would significantly simplify the shaping using compaction, hot pressing, slip casting, injection moulding, machining in green and presintered state. But, one must grant that a minimal shrinkage leads to a minimal warpage, and in the end to better accuracy. However, its application in dentistry is not known.
Machinable Glass -Ceramic (MGC) Another material-focused approach is the development of glass-ceramic materials like Macor® MGC1 or Dicor® MGC2. They are designed especially for improved machining properties in turning, milling, boring, grinding with conventional tools. Corning states for Macor good accuracy of 0.01 mm in machining. However, Macor possesses a strength lower than 150 MPa and a toughness below 1.5 MPa√m. Dicor is already applied in dental application although it exhibits similar weak mechanical properties as Macor. The high failure rates in clinical use of Dicor are unacceptable [61]. 1. 2.
Corning Inc., New York 14831, USA DENTSPLY International, York, PA 17405-0872, USA
29
State of the Art
Rapid Prototyping Rapid prototyping (RP) technology or solid freeform fabrication is known for its ability to produce complex shapes and to be suitable for small batch sizes [6, 7]. Usually RP technology starts with a CAD solid or surface model which is sliced into many layers in order to generate the tool path information for a computer controlled device. Ceramic RP adopted the methods from building polymer parts to the fabrication of ceramic green bodies. An exception was three dimensional printing which was specifically developed for ceramic material. Griffin et al. describe laminated object manufacturing (LOM) for alumina, ceria doped zirconia, and alumina - zirconia mixtures having good mechanical properties, more than 95 % theoretical density (TD), and less than 1 % open porosity [62-64]. Possible dimensional accuracy is about 0.25 mm. Not only laminated object manufacturing was applied to ceramic materials but other prominent commercialized RP techniques were used successfully, too. However, accuracy is obviously insufficient, the process requires a proper CAD model, and the long fabrication times prevent its application in dentistry at the time being.
Joining of single units for fabrication of multiple-unit bridges Recent approaches fabricate single ceramic restoration units and join them together with a glass solder [37]. The design and location of the solder joint are crucial for its success because the interdental connectors experience maximum stresses. However, this interesting technology has not yet proven its clinical feasibility for ceramic restoration.
III.9
Discussion and Summary
Fabrication of dental restorations need adequate skills, technology, materials, and devices. Skills are the knowledge and handcraft skills set by the dental technicians actually working in that field. Secondly technology defines the reliability and simplicity of the fabrication process of dental restorations, and therefore, determines the skills needed. Technology limits the shapes and accuracy of the restorations as well as the applicable materials. Thirdly, materials should exhibit several favourable properties like high strength, high toughness, tooth-like colour, high translucency, and high biocompatibili-
State of the Art
30
ty. Fourthly devices such as CAD/CAM incorporate a certain technology and support the technicians, but devices have their price, require maintenance and additional skills. The main requirements are to fabricate clinically applicable all-ceramic restorations with the required accuracy, in a reliable and fast way, at low costs out of high strengh, toughness and translucent ceramic. Conventional hard machining from prefabricated ceramic blanks gains from lot of experience gathered in the last decades. Ease and reliability of fabrication is probable. However, it requires expensive machines, long machining times, special tools to withstand cutting forces and temperatures, and in case of grinding wear dimensional control for the tools. High cutting forces limit tooling for hard machining and in consequence limit its shape fabrication capability. From the technological view hard machining suffers from chipping defects at edges, surface flaws and microcracks beneath the surface and hence, mechanical strength and reliability of the ceramic product are lowered. Especially for dental restorations these facts are not tolerable. Ultrasonic and electrodischarge machining both require to fabricate extra tools. Ultrasonic, electrodischarge, la-
ser and laser-assisted machining limit the choice of possible materials. Forming methods like compaction, slip casting, hot pressing, and injection moulding require to prefabricate moulds or forms compensating the sintering shrinkage of the material. Therefore they are complicated and expensive and in addition require the skilled handling of either a slip, a paste, or a powder. The ceramic processes behind these forming methods are considered to be very complex, even for experienced dental technicians [5]. Thus fabrication starting from the powder may be performed only under industrially conditions as in case of Procera. Shrinkage-free ceramics are at the beginning of its dental application. Machinable glass-ceramics are undergoing research but their potential is limited by their mechanical properties. Rapid prototyping has a great potential but its major drawbacks are long fabrication time, the low accuracy, and the high expenditure in devices. On the other hand, soft machining - whether green or presintered state - is said to have high potentials [2, 65, 66] despite its drawbacks which are the poor accuracy and the handling of mechanically weak green bodies. A work-over in the hard state may be required in order to achieve the necessary accuracy. The mechanical weak properties in conjunction with thin wall thickness of less than 0.5 mm and very sharp edges in case of
31
State of the Art
dental restoration may lead to high failure rates during softmachining. Moreover, experience in soft machining is empirical and not very well disseminated. In literature there are examples and materials mentioned: beryllia blocks, discs, tubes, crucibles, spark plugs insulators in near-net shape with ± 1 % relative dimensional tolerance [67], alumina monolithic ceramic cutting tools and insulators in near-net shape [68], spark plug insulators [69]. Benefits of soft machining comprise the easy fabrication of individual and complex shapes, time and cost savings apply, and no necessity for mould fabrication and wax burn-out.
III.10 References [1]
DIN: "Fertigungsverfahren. Begriffe und Einteilung. (Entwurf)". 1985, Beuth Verlag: Berlin.
[2]
König, W. and Wagemann, A.: "Machining of ceramic components: process-technological potentials." in International Conference on Machining of Advanced Ma-
terials. 1993. Gaithersburg, Maryland: Materials Science and Engineering Laboratory, National Institute of Standards and Technology. [3]
Klocke, F.: "Modern approaches for the production of ceramic components." Jour-
nal of the European Ceramic Society, 1997. 17: p. 457-465. [4]
Richerson, D.W.: "Modern Ceramic Engineering. Properties, Processing, and Use in Design.", 2nd Edition, Revised and Expanded ed. 1992, New York: Marcel Dekker, Inc.
[5]
With, G.d.: "Process Control in the Manufacture of Ceramics." in Processing of
Ceramics. Part I. , R.J. Brook, Editor. 1996, VCH Verlagsgesellschaft mbH: Weinheim, New York, Basel, Cambridge, Tokyo. p. 27-67. [6]
Lenk, R.: "Rapid Prototyping of Ceramic Parts." Advanced Engineering Materials, 2000. 2(1-2): p. 40-47.
[7]
Halloran, J.W.: "Freefrom Fabrication of Ceramics." British Ceramic Transac-
tions, 1999. 98(6): p. 299-303.
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[8]
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Becker, J. and Heidemann, D.: "Entwicklungsstand und Probleme des Einsatzes von CAD/CAM-Systemen." Deutsche Zahnärztliche Zeitung , 1993. 48(10): p. 611-617.
[9]
Rekow, E.D.: "Computer-aided design and manufacturing in dentistry: A review of the state of the art." The Journal of Prosthetic Dentistry, 1987. 58(4): p. 512516.
[10] Rekow, D.E.: "High-Technology Innovations - and Limitations - for Restorative Dentistry." Restorative Dentistry, 1993. 37(3): p. 513-524. [11] Zel, J.M.v.d.: "Heutige CAD/CAM-Systeme im Vergleich." Quintessenz Zahn-
technik, 1999. 25(2): p. 193-204. [12] Kunzelmann, K.H. and Hickel, R.: "The machinability of different dental materials for cad/cam systems." in International Conference on Machining of Advanced Ma-
terials. 1993. Gaithersburg, Maryland: Materials Science and Engineering Laboratory, National Institute of Standards and Technology. [13] Hickel, R. and Kunzelmann, K.-H.: "Keramikinlays
und Veneers.". 1997,
München, Wien: Hanser. [14] Mehl, A. and Hickel, R.: "Aktueller Entwicklungsstand und Perspektiven von maschinellen Herstellungsverfahren für Zahnrestaurationen." International Jour-
nal of Computerized Dentistry, 1999. 2(8): p. 9-35. [15] Reed, J.S.: "Principles of Ceramics Processing.", 2nd Edition ed. 1995, New york: John wiley & Sons, Inc. [16] Ivoclar: "IPS Empress.", 1997, Scientific Service, Research and Development, Ivoclar AG: Schaan, Principality of Liechtenstein. [17] Höland, W., Schweiger, M., Frank, M., and Rheinberger, V.: "A Comparison of the Microstructure and Poperties of the IPS Empress2 and the IPS Empress GlassCeramics." Journal of Biomedical Materials Research, 2000. 53: p. 297-303. [18] Dong, J.K., Lüthy, H., Wohlwend, A., and Schärer, P.: "Heat-Pressed Ceramics: Technology and Strength." The International Journal of Prosthodontics, 1992. 5(1): p. 9-16. [19] Kramer, C., McLaughlin, J.F., DeLuca, R.D., Hornor, J.A., Daub, M., and Andino, K.: "Low-Fusing Temperature Porcelain, Compositions, Prosthesis, Methods and Kits.", 1997, World Intellectual Property Organization, WO 97/30678.
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State of the Art
[20] Ivoclar: "IPS Empress 2.", 1999, Scientific Service, Research and Development, Ivoclar AG: Schaan, Principality of Liechtenstein. [21] Oh, S.-C., Dong, J.-K., Lüthy, H., and Schärer, P.: "Strength and Microstructure of IPS Empress2 Glass-Ceramic after Different Treatments." International Journal of
Prosthodontics, 2000. 13(6): p. 468-472. [22] Märki, H.-B.: "Der Additiveinfluss auf das Sinterverhalten von tetragonal stabilisiertem Zirkonoxid.", semester thesis 91/10, 1991, in ETH-Ceramics. Swiss Federal Institute of Technology: Zurich. [23] Sägesser, P.: "Einfluss von Sinteradditiven auf die Mikrostruktur von TZP-Sinterproben.", semester thesis 90/20, 1990, in ETH-Ceramics. Swiss Federal Institute of Technology: Zurich. [24] Zemp, P.: "Der Additiveinfluss auf das Sinterverhalten von tetragonal stabilisiertem Zirkonoxid.", semester thesis 91/03, 1991, in ETH-Ceramics. Swiss Federal Institute of Technology: Zurich. [25] Cantz, T.: "Entwicklung von Heipressmatrizenmaterials zur Herstellung
von
Dentalkeramik.", semester thesis 91/07, 1991, in ETH-Ceramics . Swiss Federal Institute of Technology: Zurich. [26] Ha, S.-W. and Riedi, M.: "Untersuchung von Feuerfestmaterialien hinsichtlich ihrer Eignung als Heisspressmatrizenmaterial.", semester thesis 91/02, 1991, in
ETH-Ceramics. Swiss Federal Institute of Technology: Zurich. [27] Menet, M.: "Entwicklung von Feuerfestmaterial zum Heisspressen von Dentalkeramik.", diploma thesis 92/05, 1992, in ETH-Ceramics. Swiss Federal Institute of Technology: Zurich. [28] Grossman, D.G. and Johnson, J.L.M.: "Glass-Ceramic Compositions for Dental Constructs.", 1987, United States Patent, US 4,652,312. [29] Fries, R. and Rand, B.: "Slip Casting and Filter-Pressing." in Processing of Ce-
ramics. Part I., R.J. Brook, Editor. 1996, VCH Verlagsgesellschaft mbH: Weinheim, New York, Basel, Cambridge, Tokyo. p. 153- 187. [30] Claus, H.: "Vita - In-Ceram, ein neues Verfahren zur Herstellung oxidkeramischer Gerüste für Kronen und Brücken." Quintessenz Zahntechnik, 1990. 16(1): p. 3546.
State of the Art
34
[31] Claus, H.: "The structure and microstructure of dental porcelain in relationship to the firing conditions." Int J Prosthodont, 1989. 2(4): p. 376-384. [32] Kappert, H.F., Knode, H., and Manzotti, L.: "Metallfreie Brücken für den Seitenzahnbereich. Herstellungsverfahren und Festigkeit von In-Ceram-VollkeramikBrücken." dental-labor, 1990. 37(2): p. 177-183. [33] Wolf, U.: "Das Wolceram-Verfahren - der neue Weg in der In-Ceram-Technik."
Quintessenz Zahntechnik, 1999. 25(3): p. 289-294. [34] Tissot, J.: "Schwindungsfreie Herstellung einer zirkonoxidhaltigen Keramik.", diploma thesis 90/05, 1990, in ETH-Ceramics. Swiss Federal Institute of Technology: Zurich. [35] Andersson, M. and Odéon, A.: "A New All-ceramic Crown: A Dense-Sintered, High Purity Aluminia Coping with Porcelain." Acta Odontol Scand, 1993. 51: p. 59-64. [36] Oden, A., Andersson, M., and Antonson, I.: "Method of Manufacturing Ceramic Tooth Restorations.", 1994, World Intellectual Property Organization, WO 94/ 27517. [37] Oden, A. and Salomonson, J.: "Method of Manufacturing Ceramic Artificial Tooth Bridges.", 1999, World Intellectual Property Organization, WO 99/13795 (EP 1011512). [38] Hegenbarth, E.A.: "Procera Aluminium Oxide Ceramics: A new way to achive stability, precision and esthetics in all-ceramic restorations." QDT, 1996: p. 21-34. [39] Oden, A., Andersson, M., Krystek-Ondracek, I., and Magnusson, D.: "A 5-year Clinical Follow-up Study of Procera AllCeram crowns." Journal of Prosthetic
Dentistry, 1998. 80(4): p. 450-455. [40] Enlund, B.: "Nobel Biocare Interim Report 3. January-September 2000.", 2000, Ernst & Young AB, Gothenburg, Authorized Public Accountant: Gothenburg. [41] Reckziegel, A. and Willmann, G.: "Konstruieren mit Keramik - Herstellung und Masshaltigkeit ohne Nachbearbeitung." Sprechsaal, 1985. 118(4): p. 332-338. [42] Warnecke, G., Rosenberger, U., and Mohr, H.: "Aufgabengerechte System- und Prozessauslegung beim Schleifen von Hochleistungskeramik." in Ergebnispräsen-
tation des BMFT-Verbundprojektes: Schleifen von Hochleistungskeramik., G. Warnecke, Editor. 1994, Universität Kaiserslautern: Kaiserslautern. p. 18.1-18.12.
35
State of the Art
[43] Jahanmir, S., Xu, H.K., and Ives, L.K. ( eds.): "Mechanisms of Material Removal in Abrasive Machining.", Machining of Ceramics and Composites. In: Manufacturing Engineering and Materials Processing., ed. S. Jahanmir, M. Ramulu, and P. Koshy. Vol. 53. 1999, Marcel Dekker: New York, Basel. 11-84. [44] Rostvall, T.: "Method and Device for the Abrasive Presicion Machining of a Blank.", 1998, World Intellectual Property Organization, WO 98/36871. [45] Klocke, F., Verlemann, E., and Schippers, C. ( eds.): "High-Speed Grinding of Ceramics.", Machining of Ceramics and Composites. In: Manufacturing Engineering and Materials Processing., ed. S. Jahanmir, M. Ramulu, and P. Koshy. Vol. 53. 1999, Marcel Dekker: New York, Basel. 119-137. [46] Kosmac, T., Oblak, C., Jevnikar, P., Funduk, N., and Marion, L.: "Strength and Reliability of Surface Treated Y-TZP Dental Ceramics." Journal of Biomedical
Materials Research, 2000. 53: p. 304-313. [47] Sindel, J., Petschelt, A., Grellner, F., Dierken, C., and Greil, P.: "Evaluation of Subsurface Damage in CAD/CAM Machined Dental Ceramics." Journal of Mate-
rials Science: Materials in Medicine, 1998(9): p. 291-295. [48] Dörfel, I. and Urban, I.: "TEM-Untersuchungen der Randzonenschädigung von Aluminiumoxidkeramiken nach Schleifbearbeitung." Mat.-wiss. u. Werkstofftech., 1996. 27: p. 307-312. [49] Pfeiffer, W. and Hollstein, T.: "Characterisation and Assessment of Machined Ceramic Surfaces." Interceram, 1997. 46(2): p. 98-105. [50] Tönshoff, H.K., Wobker, H.G., and Lierse, T.: "Randzoneneigenschaften und Oberflächenausbildung geschliffener Aluminiumoxidkeramik." Mat.-wiss. u.
Werkstofftech., 1996. 27: p. 323-330. [51] Zhang, G.M., Anand, D.K., Ghosh, S., and Ko, W.F.: "Study of the Formation of Macro- and Microcracks During Machining of Ceramics." in International Con-
ference on Machining of Advanced Materials. 1993. Gaithersburg, Maryland: Materials Science and Engineering Laboratory, National Institute of Standards and Technology. [52] Mikrona Technologie ( ed. "Celay: directe oder indirekte Herstellung keramischer Formkörper. Bedienungsanleitung." 11/93 ed,, ed. M.T. AG. 1993, Mikrona Technologie AG: Spreitenbach.
State of the Art
36
[53] Spur, G.: "Keramikberarbeitung: Schleifen, Honen, Läppen, Abtragen.". 1989, München, Wien: Carl Hanser Verlag. [54] Pröbster, L., Groten, M., and Girthofer, S.: "Kopiergefräste, glasinfiltrierte Aluminiumoxid-Keramikkronen: Das Celay-In-Ceram-System step by step." Phillip
Journal, 1994. 11(12). [55] Klocke, F. and Hilleke, M. ( eds.): "Ultrasonic Machining of Ceramics.", Machining of Ceramics and Composites. In: Manufacturing Engineering and Materials Processing., ed. S. Jahanmir, M. Ramulu, and P. Koshy. Vol. 53. 1999, Marcel Dekker: New York, Basel. 483-524. [56] Hahn, R. and Schulze, P.: "Sonoerosive Fertigung komplexer keramischer Kleinserienbauteile." cfi/Ber. DKG, 1993. 70(7). [57] Andersson, M., Bergman, B., Bessing, C., Ericson, G., Lundquist, P., and Nilson, H.: "Clinical results with titanium crowns fabricated with machine duplication and spark erosion." ActaOdontol. Scand., 1989. 47: p. 279-286. [58] Klocke, F. and Zaboklicki, A. ( eds.): "Laser-Assisted Turning of Ceramics.", Machining of Ceramics and Composites. In: Manufacturing Engineering and Materials Processing., ed. S. Jahanmir, M. Ramulu, and P. Koshy. Vol. 53. 1999, Marcel Dekker: New York, Basel. 551-574. [59] Lee, S.-J., Kriven, W.M., and Kim, H.-M.: "Shrinkage-Free, Alumina-Glass Dental Composites via Aluminum Oxidation." Journal of the American Ceramic Soci-
ety, 1997. 80(8): p. 2141-2147. [60] Hennige, V.D., Hausselt, J., Ritzhaupt-Kleissl, H.-J., and Windmann, T.: "Shrinkage-free ZrSiO4-ceramics: characterisation and applications." Journal of the Euro-
pean Ceramic Society, 1999(19): p. 2901-08. [61] Felden, A., Schmalz, G., Federlin, M., and Hiller, K.-A.: "Retrospective Clinical Investigation and Survival Analysis on Ceramic Inlays and Partial Ceramic Crowns: Results up to 7 Years." Clinical Oral Investigation, 1998. 2: p. 161-167. [62] Griffin, A., Mumm, D.R., and Marshall, D.B.: "Rapid Prototyping of Functional Ceramic Composites." The American Ceramic Society Bulletin, 1996. 75(7): p. 6568. [63] Griffin, A.: "Rapid Prototyping with Engineered Ceramics." Ceramic Industry, 1997(April): p. 86-88.
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State of the Art
[64] Griffin, C., Daufenbach, J.D., and McMillin, S.: "Desktop Manufacturing: LOM vs Pressing." The American Ceramic Society Bulletin, 1994. 73(8): p. 109-113. [65] Klocke, F.: "Modern Approaches for the Production of Ceramic Components. Extended Abstract." in International Conference on Shaping of Advanced Ceramics. 1995. Mol, Belgium: VITO, Boeretang 200, 2400 Mol, Belgium. [66] Klocke, F., Gerent, O., and Schippers, C.: "Green Machining - Economical and Technical Potentials." Production Engineering, 1996. III(1): p. 15-18. [67] Kovell, G.A., Sepulveda, J.L., and Wellman, B.: "Optimization Green Machining after Isopressing of Beryllia Ceramic Bodies." in Forming Science and Technolo-
gy for Ceramics., M.J. Cima, Editor. 1991, The American society of Ceramics: Cincinnati, Ohio. p. 231-239. [68] Halcomb, D.L. and Rey, M.C.: "Ceramic cutting tools for machining unsintered compacts of oxide ceramics." Ceramic Bulletin, 1982. 61(12): p. 1311-1314. [69] Sheppard, L.M.: "The challenges of ceramic machining continue." The American
Ceramic Society Bulletin, 1992. 71(11): p. 590-610.
State of the Art
38
39
IV.
Direct Ceramic Machining Process
All-Ceramic Teeth Restorations: The Direct Ceramic Machining (DCM)-Process Abstract
The Direct Ceramic Machining (DCM) process allows to fabricate all-ceramic multiple teeth restorations. The idea is to machine the load bearing framework of a restoration from a soft, porous ceramic zirconia blank in an enlarged shape. The individual shape of a restoration is digitally acquired from a positive model of the framework. Shrinkage to the desired dimensions avoids any hard machining. Then, the framework is coated with veneer porcelain in order to meet individual aesthetic appearance before finally cementing the restoration in the patient’s mouth. In this chapter the fabrication of the blank, its coloration and its properties are described. Bilayered structures of dense sintered zirconia with a veneering porcelain adjusted in thermal expansion coefficient to zirconia were mechanically tested. Fabrication steps, e.g. digitizing, milling, sintering and veneering of multiple unit bridges using prefabricated blanks according to the DCM process are discussed. Preclinical performance of three-unit frameworks and test bridges was evaluated using bend testing. Finite element analysis of three-unit frameworks was performed. Results of zirconia bridges show that they have the highest strength in combination with the highest reliability compared to other commercially available dental materials systems. The clinical application of zirconia bridges demonstrate their performance as single or multipleunit restorations for the posterior region which requires high load bearing capacity as well as for anterior region requiring aesthetic appearance.
Direct Ceramic Machining Process
40
IV.1
Ceramic Blank: Powder, Fabrication, and Properties
IV.1.1
Introduction
Fabrication of good ceramic components starts with the selection of the raw materials, in our case this was a ceramic powder consisting of 3 mol-% yttria-stabilized zirconia. Blanks that are prepared from this powder are an intermediate product used by dental technicians for dental restorations. Blank fabrication for DCM is one of the crucial process steps as it determines the accuracy, strength and reliability of the fabricated dental restoration. The blanks require-
ments are the isotropic shrinkage in all spatial directions, the ability to sinter to full density, the easy machinability, and the freedom from defects. The first prerequisite means the blank to have a homogeneity of particle and pore distribution on a length scale of a few micrometers to the whole size of the blank. Second prerequisite concerns homogeneity on the length scales of a few particle diameters (up to 200 nm) in order to eliminate all pores in the sintered part at low sintering temperatures to achieve a high strength combined with a high reliability [1]. The third prerequisite of easy machinability is achieved by weakly bonded ceramic particles on a length scale of a few nanos up to a few microns. The fourth prerequisite is to have reliable defect-free blanks, which means no flaws in the blanks and later in the sintered component to meet International Standards Organization (ISO) regulations [2]. Choosing appropriate and economic ceramic forming method for blanks fabrica-
tion depends on the amount and size needed for the blanks. More than 10 millions tooth units were inserted in 1996 in Germany [3]. Therefore high volume forming methods enabling the economic production of more than 100’000 blanks per year must be chosen. Each restoration is composed of one or more tooth units and therefore requires an appropriate sized blank. For example, a mandibular posterior five-unit bridge from the first premolar to the third molar shows an anatomical length of 46 mm [4] which determines
41
Direct Ceramic Machining Process
the blanks’ length. Cranio-caudal and buccal-palatinal curvature of the denture determine the blanks’ height and width. Hence, blank sizes up to 25 mm x 25 mm x 70 mm are needed based on the anatomical dimensions, taking into account the approximated shrinkage of 25 %, and allowing some excess material at each font side of the blank. Conventional forming methods for ceramic porous blanks are uniaxial or isostatic compaction, extrusion, slip or DCC casting, injection moulding [5-7]. Selected characteristics for these forming methods are compiled in tab. 1. Compaction uses commercially available dry powders for consolidation and shaping of the green part in one step. Die compaction uses rigid moulds under uniaxial or multiaxial pressure [6, 8]. Stages of die compaction include filling of the die with powder, compaction and the ejection. Homogeneity e.g. density and porosity varies within the compact due to particle-particle, particle-die wall and particle-punch friction. Density variations are reported in the range of 3 to 10 % [7, 9-11]. Die compaction is often applied for mass production because of its simplicity, speed and automation. Parts range widely in size from a few mm to 10 cm. Cold Isostatic Pressing (CIP) uses soft plastic or rubber moulds under isostatic pressure [6]. Therefore, homogeneity is superior to die compaction, and density variations are reported to be in the range of 1 to 2 % [7, 12]. CIP is also applied in mass production but its productivity is lower compared to die compaction due to lower speed. Parts may be sized from few mm to 0.5 m. The spark plug is one of the best known examples [5]. Slip casting uses a suspension consisting of ceramic particles in a liquid dispersion medium [7]. Porous moulds extract the liquid and consolidate the suspension at the walls. Essentials for slip casting are the control of the pH of the slurry, its deflocculation and its rheology. Parts may range from a few mm to several meters. Slip casting with fugitive-wax forms enables the fabrication of very complex shapes like gas-turbine rotors or very simple shapes like plates or crucibles. Direct Coagulation Casting (DCC) uses an electrostatically stabilized, low viscosity aqueous suspension of ceramic particles [13]. Consolidation is performed by time-delayed reactions either shifting pH or changing ionic strength leading to coagulation. DCC produces parts of high homogeneity which may range from a few mm to 0.2 m. Industrial exploitation currently is under realization. Injection moulding and extrusion use a mixture of powder and up to 50 vol-% additives that plastically deforms under pressure [7]. Extrusion forces the pasteous mixture through a shaped die. It is applied in mass production and can be performed continuous-
Direct Ceramic Machining Process
42
ly using auger extruders or discontinuously in batches using piston extruders. Parts with a constant cross-section such as bricks, insulators, heat-exchangers are continuously extruded. Injection moulding forces the plastic mixture under high pressure through a nozzle in a closed mould. Accuracy and productivity of injection moulding is high. Parts vary in size from a few mm to more than 100 mm and in shape complexity from simple like seal rings to highly sophisticated and complex shaped turbo charger rotors [14]. Wet forming methods need special knowledge in handling the suspension and an additional drying step under controlled conditions before sintering. High amounts of organic additives in pastes require a sophisticated debinding process and may decrease homogeneity. DCC is one potential candidate for blank production, its industrial application is currently under development. Dry forming methods like die and isostatic pressing use ready-to-go powders requiring no special handling knowledge. However, without further machining of the compacted components these parts do not fulfil the dimensional tolerance required for DCM [15]. Therefore considering all the various aspects CIP was used as forming method for blank fabrication in this work with an additional trimming step.
43
Direct Ceramic Machining Process
Forming Method
Feed
Homogeneity
Tolerance
Productivity
Cold Isostatic Pressing
powder
1 -2 %
± 0.5 % (± 3 %)
high, stroke rates of up to 2 s-1 per green body very high, stroke rates up to 100 s-1 per green body
die pressing
powder
3 - 10 %
± 0.5 % (± 1-3 %)
extrusion
paste
bad, high amount of binder
± 1.5 % (± 3-5 %)
low, time for drying and debinding
slip casting
suspension
good
± 0.5 % (± 3-5 %)
low, time for drying
Direct Coagulation Casting
suspension
good
-
high, time for drying
injection moulding
paste
bad, high amount of binder
± 1.5 % (± 3 %)
low, allow enough time for drying and debinding
Tab. 1:
Comparison of ceramic forming methods enabling a fabrication of blanks. Dimensional tolerance values in brackets are valid using precision processes without any hardmachining afterwards [15].
Fabrication of ready-to-use blanks consists of the fabrication steps shown in fig. 3. CIP is used as the initial forming method for consolidating and preshaping the powder to an appropriate green body. Then, the green body is trimmed to accurate shape and dimension and afterwards heat treated for debinding and adjusting machinability at the same time. Presintered density is determined in order to calculate the appropriate shrinkage-compensating enlargement factor for the specific blank. Colouring of the blank by aqueous solution infiltration may follow afterwards.
Fig. 3:
chapter IV.1.9
Fabrication Process
Direct Ceramic Machining Process
44
Powder Powder
chapter IV.1.2
Cold Cold Isostatic Isostatic Pressing
chapter IV.1.3
Trimming Trimming
chapter IV.1.4
Green GreenBody Body
chapter IV.1.5
Heat Heat Treating Treating
chapter IV.1.6
Blank Blank
chapter IV.1.7
Colouring Blank Colorating
chapter IV.1.8
Steps for the fabrication of blanks.
The following chapters will focus on the fabrication of blanks for dental application. The complete blank fabrication will be discussed starting from the powder and ending with the blanks characteristics as functions of compaction pressure and heat treatment temperature. Colouring of the blanks is desirable for some clinical applications for a better aesthetic appearance of the dental restoration. Colours may be introduced in the powder state or / and in the presintered blank. In this work the colouring of the blanks has been investigated too. Characteristics of the fabrication process as found in this work with laboratory equipment may provide knowledge about the tolerances helpful for setting up industrial production.
45
IV.1.2
Direct Ceramic Machining Process
Powder
Zirconia powders (tetragonal zirconia polycrystals (TZP) stabilized with 3 mol-% yttria) from different suppliers were compared with respect to powder characteristics, green body properties, sintering behaviour and microstructure of sintered bodies, see [16-18]. The TZ-3YB grade from Tosoh Corporation (Tokyo, Japan) is commercially available in large quantities and fulfils the basic requirements. Biocompatibility was established by histological analysis [19, 20], and conformity with ISO-norms [2, 21, 22] was already investigated earlier. This powder is already applied for hip joint femoral heads [23-27], dental root canal posts for devital retainer teeth [28-31], and abutments for dental implants [32, 33]. Powders of TZ-3YB grade and TZ-3YB-E grade were used in this study. The TZ-3YB and the TZ-3YB-E grade are powders with acrylic binder additive which makes them especially suitable for pressing. Characteristic chemical data of the TZ-3YB powder is shown in tab. 2. Typically ZrO2 powder contains approximately 2 wt-% HfO2 which is difficult to separate from the zirconia due to its chemical similarity. Therefore, HfO2 is included in the amount of ZrO2. The variance in chemical composition of different TZ-3YB powder lots is of minor magnitude (appendix: chapter VI.1,
tab. 17). For TZ-3YB-E the chemical analysis is listed in chapter VI.1, tab. 18 (appendix).
Direct Ceramic Machining Process
Tab. 2:
46
contents
average amount [wt-%]
variance [%]
ZrO2 (+ HfO2)
94.860 ± 0.061
0.13
Y2O3
5.108 ± 0.060
2.35
Al2O3
< 0.005
-
SiO2
0.005 ± 0.004
-
Fe2O3
0.002 ± 0.001
-
Na2O
0.020 ± 0.004
-
Ignition Loss
3.443 ± 0.160
9.30
Average chemical composition of TZ-3YB and its variance. Average data was calculated from 8 different TZ-3YB powder lots (according to appendix chapter VI, tab. 17).
Powder and materials characteristics are compiled in tab. 3. The TZ-3YB powder consists of spray dried nearly ideal spherical granules. They are shown in fig. 4 and exhibit a mean 60 µm diameter. The granules consist of smaller particles so called agglomerates of 0.32 µm median diameter (D50 ). Its size distribution measured by dynamic light scattering method1 is presented in fig. 5. However, median values as analysed here are half of the values of the Tosoh’s specification as contained in tab. 3 which were obtained by ultra sedimentation method. Differences may be due to using different analysis methods. The agglomerates are build up of much smaller primary crystals. These primary crystal are of nanometer size, and in the order of a thousandth of the powder granule diameter. However, primary crystal size in transmission electron microscopy (TEM) shows a 40 to 60 nm diameter whereas Tosoh’s specification based on X-ray diffractometer analysis refers to 27 nm diameter. This also may be due to different methods. The fluidity of the powder is a crucial characteristic for a homogeneous filling of the press moulds [17]. Fluidity of both powders is guaranteed by the spherical shape of agglomerates granules and is important for production control.
1.
Ultrafine Particle Analyser (UPA), Microtrac, Honeywell (now: Nikkiso Co., Tokyo, Japan)
47
Direct Ceramic Machining Process
To summarize, the material properties of the sintered parts prove TZ-3YB-E being superior in bend strength to TZ-3YB and possessing the smaller grained microstructure. Toughness, hardness, sintered density are similar for both materials. However, sintering temperature to achieve full density is 1500 °C for TZ-3YB and 1320 °C for TZ-3YB-E, respectively. TZ-3YB-E contains small amounts of alumina in contrast to TZ-3YB. TZ3YB has already being commercialized by Tosoh since years whereas the TZ-3YB-E has been launched during this work. Therefore, there is an advantage of experience in the processing and the behaviour of the TZ-3YB powder. Both powders may be suitable however for use in dental restorations.
powder/Beatrix_Michel/ powder_primar_grain.jpg
Fig. 4:
powder/TZP-03_100um_granule.jpg
Primary particles and spray dried granules of TZ-3YB. Left: Primary particles in TEM [18] showing a size of 40 to 60 nm. Right: Spherical spray granules in light microscopy showing a size of 60 µm [34].
Direct Ceramic Machining Process
property
unit
fluidity
48
TZ-3YB
TZ-3YB-E
value
reference
g/s
0.24
[17]
specific surface area
m2/g
16
[34]
bulk density
g/cm3
1.1
primary crystal size (TEM)
nm
27
[34]
27
[35]
agglomerate size (median, D50)
µm
0.6
[34]
0.6
[35]
spray granule size (median, D50)
µm
60
[34]
60
[35]
green density (die pressed, 69 MPa)
g/cm3
2.66
[34]
2.66
[35]
green density (CIP, 300 MPa)
g/cm3
3.19
green strength (die pressed, 69 MPa)
MPa
20
[36]
sintered density (69 MPa / 1500 °C / 2 h)
g/cm3
6.05
[35]
grain size (REM) (300 MPa / Ts/ 2 h)
nm
322 (1500 °C)
bend strength
MPa
1000
[35]
1100
[35]
toughness
MPa√m
5
[34]
5
[35]
hardness
HV10
1250
[34]
1250
[35]
Tab. 3:
value
reference
16
[35]
1.1
3.19
6.05
[35]
150 (1320 °C)
TZ-3YB and TZ-3YB-E powder and sintered part characteristics.
49
Direct Ceramic Machining Process
80
6
60
Kalleidagraph-Data: powder/TZ-3YB.qdc 4 Image: powder/particlesize_TZ-3YB.wmf
40
2
20
0
0.32 µm
Counts [%]
8
0.1
1
Cumulative [%]
100
10
(a)
0
Counts [%]
(b)
30
100
25
80
Kalleidagraph-Data: powder/E-Grade.qdc 20 Image: powder/particlesize_e-grade.wmf 15
40 0.24 µm
10 5 0
60
0.1
Cumulative [%]
Agglomerate Size [µm]
20 1
0
Agglomerate Size [µm] Fig. 5:
Monomodal agglomerate size distribution for TZ-3YB (a) and TZ-3YB-E (b) powder analysed by dynamic light scattering method. Median agglomerate size (D50) is 0.32 µm for TZ-3YB and 0.24 µm for TZ-3YB-E, respectively.
Direct Ceramic Machining Process
IV.1.3
50
Cold Isostatic Pressing (CIP)
Prior to compaction an appropriate amount of TZ-3YB powder was filled in a sealed elastic containment which was made of plastic material1 and which had a cylindrical shape. For different sizes of blanks containments with different heights and diameters were used. A wall thickness of about 5 mm was necessary for good stiffness of the containment in order to get good cylindrical shape of the pressed green bodies. 90 g of powder were used for a mould of approximately 38 mm diameter and 75 mm height. Special attention was paid to the filling step, because this step is a source of error which propagates through the complete DCM process. For fabrication of high quality blanks containments had to be dust clean and freed from agglomerates after each pressing. The surrounding was clean in order to avoid contamination of the mould, and a steady flow of powder from a glass beaker into the containment was ensured in order to get a homogeneous filling. No tapping or riddling was used to get a higher initial filling density. The mould was closed with a conical top of equal elasticity and then packed in two tied up elastic bags2 to prevent any contamination of the inner powder by the pressure fluid. The containment was then placed in the pressure vessel. Pressure was applied slowly up to a maximum of 300 MPa with a dwell time of two minutes3. Compaction pressure was reduced gently to atmospheric pressure because abrupt pressure release caused cracking of the brittle green body. Then the containment was removed from the pressure vessel and unwrapped from the elastic bags. The green body was easily taken out of the mould because the consolidated powder compact is significantly smaller than the mould. Any contamination of the green body with pressure fluid was avoided during unpacking procedure. Atmospheric humidity causes uncontrollable swelling of the binder in the green body which therefore had to be stored in a dry drying oven4 at 120 °C. The materials used for pressing and the resulting green bodies are shown in fig. 6. For cylinders 90 g powder resulted in a green body of about 26 mm diameter and approximately 52 mm height.
1. 2. 3. 4.
Plastisol 40° Shore, EPSI, Temse, Belgium Disposan AG, Würenlos, Switzerland KIP 100 E, Paul-Weber Maschinen- und Apparatebau, Remshalden-Grunbach, Germany WU6100, Heraeus Instruments, Zürich, Switzerland
51
(a)
Direct Ceramic Machining Process
(b)
figures/raw_material_007.jpg
Fig. 6:
figures/raw_material_008.jpg
(a) Materials for wet-bag-CIP: powder, plastic form, top, elastic bags, unpacked green body. (b) Various sizes and shapes of green bodies produced, according to the end part dimensions. White arrows: elephant feet.
IV.1.4
Trimming of the Green Body
Depending on the outer shape of the blanks which is either cylindrical or cubic the green bodies were ground and cut using a flat grinding machine1, or they were turned using a lathe2. In case of cubic shapes machining in the presintered state was preferred because grinding required gluing bodies on a steel plate. Cylindrical shape was preferred because of the simple fabrication process compared to the cubic shape. However, for either shape a proper geometry with undamaged edges was required in order to determine the density of the blank satisfied. For cylindrical shape high-speed-steel (HSS)3 as well as polycrystaline diamond (PCD) 4 tools were successfully applied. In case of HSS five to ten cylinders were manufactured until the tool was blunt and the surface of the cylinders showed pittings and grooves. This is an acceptable wear rate for HSS tools. For PCD tools wear rate is much
1. 2. 3. 4.
Chevalier FSG-818 AD, Falcon Machine Tools Co., Ltd., Taichung, Taiwan 102-VM, Schäublin SA, Bévilard, Switzerland Requirements for HSS tools: radius greater than 0.5 mm, clearance positive, tool orthogonal rake positive, tool cutting edge angle was chosen to 90°. Diapact (PKD) CCMW 060204 and DCMW 070204, Weiss AG, Walzenhausen, Switzerland
Direct Ceramic Machining Process
52
lower than for HSS, and more than twenty cylinders were manufactured until pittings were observed at the surface of the cylinders. Coated tools did not work at all because the coating was worn within the first seconds. A feed rate of 75 mm/min at 1100 rpm and a cutting depth of up to 2 mm were chosen. For fabricating cubic shapes the approximately cylinder shaped presintered body was cut manually on a diamond saw to have one flat surface for gluing onto a steel plate. Shellac1 with a working temperature of 120 °C was used as glue which burns out of the body with no residue below the presintering temperature. For grinding the cubic body diamond wheels for flatting2 and for cutting3 the top surface are used with a cooling agent4. For flattening a cutting depth of 0.01 mm, feed rate of 10 m/min, wheel rotation speed of 21.4 m/s, and lateral feed of 2 mm/stroke was applied. Cutting was performed in two orthogonal directions with cutting depth of 0.15 mm/stroke, feed rate of 1 m/min, wheel rotation speed of 21.4 m/s and no lateral feed. Afterwards the green bodies were detached, and glued again on the flat side before finishing the last side. Finally the shellac burn-out was performed in an electrical furnace5 at 600 °C for 4 hours with a heating rate of 1K/min.
IV.1.5
Properties of the Green Body
The average overall green body density for TZ-3YB as a function of the compaction pressure was evaluated using die pressed pellets of 11 mm diameter and 3 mm height. The relative error was estimated to 1.5 % for the density and to 2 % for the compaction pressure, respectively (appendix: chapter VI.3). Density as calculated from dimensions and mass is shown in fig. 7. Increasing compaction pressure leads to increasing overall density of the green bodies. More than 50 %TD was achieved for compaction pressures exceeding 250 MPa. Granule deformation and densification predominates for
1. 2. 3. 4. 5.
Stettler, Lyss, Switzerland Diametal D54, concentration 75, binding B2, diameter 200 mm, width 9.0 mm, Diametal AG, Biel/Bienne, Switzerland, Diametal D126, concentration 75, binding B55, diameter 200 mm, width 1.5 mm, Diametal AG, Biel/Bienne, Switzerland, Grindex SC, Blaser Swisslube AG, Hasle-Rüegsau, Switzerland N7H, Tony Güller AG, Hägendorf, Switzerland
53
Direct Ceramic Machining Process
the applied compaction pressures being sufficient for sintering to full density [6, 37]. Density as measured here corresponds to technical data sheet given by the powder supplier [36] and to the literature [17, 38]. However, CIP achieves a slightly higher green density compared to die pressing.
3.2
Curve fit: ρ = 1.3351 + 0.7087 log p R = 0.9996
sinter-properties/density/ Compaction_Density.wmf
Fig. 7:
46 43
2.6 2.4 10
53 50
3.0 2.8
56
Rel. Density [%TD]
Density ρ [g/cm3]
3.4
100
Compaction Pressure p [MPa]
40 1000
Density of TZ-3YB green bodies as function of the compaction pressure. Linear logarithmic curve fit as proposed by Reed [6] corresponds well to the measured values.
The porosity was analysed using mercury intrusion porosimetry (appendix: chap-
ter VI.3) for different compaction pressures in terms of median pore radius and pore size distribution. For the measurements cylindrical specimens of 11 mm diameter and 3 mm height were prepared for different compression pressures. Pieces from the middle of the specimens of about 0.5 g were measured. Pore size as function of the mercury cumulative volume is shown in fig. 8. For large pore sizes (right side of fig. 8) the cumulative volume shows a low remaining constant amount. Then for pore radii in the range of 30 to 20 nm cumulative volume increase quickly and reach a constant value for smaller pore radii. Therefore pore sizes have a narrow distribution from 20 to 30 nm as shown in
Direct Ceramic Machining Process
54
fig. 9. Increasing the compaction pressure shifts pore radii towards smaller values and decreases cumulative pore volume. The latter is equal to increasing density which correlates to results of fig. 7. The narrow pore size distribution in the green body is a strong indication for its homogeneity. The mean pore radius decreases for increasing compaction pressure below 300 MPa whereas a constant mean pore radius is observed for compaction pressures higher than 300 MPa. For 300 MPa compaction pressure the 20 nm mean pore radius compares very well to the 17 nm found by Chen and Mayo [38]. Porosity smaller than a 0.5 ratio of pore diameter to mean particle diameter was found to be totally eliminated during final sintering [1]. For both powders this ratio is clearly below 0.5 and therefore it is probable to achieve full density during sintering. Furthermore the shape of the cumulative-intrusion-volume curve shows a monomodal pore size distribution and therefore in combination with the magnitude of the mean porosity gives a strong indication for having closed intergranular porosity and no interfaces between the granules [6, 7]. Uniform green bodies as fabricated here are giving a good precondition for sintering to full density.
Cumulative Volume [mm3/g]
250 Compaction Pressure 150 MPa 300 MPa 500 MPa
200 150 100
sinter-properties/porosity/file obj/ Gruenkoerper_verschiedene_Pressdruecke.wmf
50 0
0.01
0.1
1
10
Pore Radius [µm] Fig. 8:
Pore radius distribution by mercury intrusion porosimetry in the TZ-3YB green body for different compaction pressures.
Counts [Arbitrary Units]
55
Direct Ceramic Machining Process
20 nm
sinter-properties/porosity/file obj/ derivative_of_cumulative_volume.WMF
0.01
0.1
1
10
Pore Radius [µm] Fig. 9:
Pore size distribution and mean pore radius for a TZ-3YB green body compacted with 300 MPa. First derivation of the cumulative volume as shown in fig. 8.
Mean Pore Radius [µm]
0.030
0.025
sinter-properties/porosity/file obj/ mean_pore_radius.WMF
0.020
0.015
0.010
0
100
200
300
400
500
Compaction Pressure [MPa] Fig. 10: Mean pore radius as function of the compaction pressure based on the mercury intrusion porosimetry.
Direct Ceramic Machining Process
56
Hardness as function of the compaction pressure was analysed using the indenter method1 with the appropriate loads (appendix: chapter VI.3). Pellets with 11 mm diameter and 3 mm height were uniaxially compacted using TZ-3YB powder. Vickers hardness as function of compaction pressure is shown in fig. 11. Hardness increases linearly as compaction pressure and density increase. No appropriate hardness data were found in literature.
Hardness [MPa]
20
15
sinter-properties/hardness/ hardness_HV10.wmf
10
5
0
0
100
200
300
400
500
Compaction Pressure [MPa] Fig. 11: Vickers hardness as a function of the compaction pressure. Analysis was performed on TZ-3YB die compacts using Vickers indenter method. Relative error of hardness is estimated to be smaller than 2 %.
1.
Zwick type 3212001/00 equipped with diamond pyramid, Zwick, August-Nagel-Strasse 5, 89079 Ulm, Germany
57
IV.1.6
Direct Ceramic Machining Process
Heat Treatment
After cold isostatic pressing and trimming of the green bodies they were heat treated in order to produce the porous blanks. The heat treatment process is characterized by its maximum temperature, heating rate and dwell time. During heat treatment the binder, water content, and absorbed moisture in the green body is removed [6]. The remainder is a binder-free blank. Chemical reactions were analysed by differential thermal analysis (DTA), thermogravimety (TG), and dilatometry. The mechanical properties of the blanks can be adjusted by the heat treatment. The important properties of the blanks in terms of density, porosity, hardness and machinability will be discussed in the following chapters. Different maximum temperatures were chosen ranging from 850 °C to 1075 °C using an electrical furnace1 and a heating rate of maximum 1 K/min. The dwell time was two hours and the samples were furnace cooled below 100 °C. In order to get a homogeneous temperature distribution in the blanks they were shielded against the electrical heating elements with refractory plates. Temperature was measured at the blanks using a thermocouple type B.
DTA / TG2 was performed with TZ-3YB and TZ-3YB-E using a heating rate of 1 K/min up to 1000 °C in air with 0.15 g of powder. The results are shown in fig. 12. Exothermic peaks were observed at 120 °C and at 300 °C. The weight loss up to 350 °C was about 2 % which compares to the data of 3.2 to 3.5 wt-% ignition loss given by the manufacturer. These results are attributed to the binder burn-out in the temperature range from 120 °C to 350 °C rather than water or moisture evaporation. Both powders show similar qualitative DTA / TG behaviour. For achieving a binder-free blank which allows sintering with high heating rates a presintering treatment at a minimum temperature of 350 °C is necessary.
1. 2.
N7H, Tony Güller AG, Hägendorf, Switzerland STA 501, Bähr Thermoanalyse GmbH, 32609 Hüllhorst, Germany
Direct Ceramic Machining Process
58
exothermic endothermic
DTA [µV]
0 3
TZ-3YB
2
TZ-3YB-E
-2 -4
TG [%]
2
4
-6 Origin Data: dta_tg/DTA_TG.opj Image: dta_tg/dta_tg.wmf
1
TZ-3YB
0 TZ-3YB-E
-1
0
200
400
600
800
1000
Temperature [°C] Fig. 12: DTA / TG for TZ-3YB and TZ-3YB-E in air using a 1 K/min heating rate.
Dilatometer1 experiments were carried out on compacts in order to follow the effects of binder removal on the dimensional stability. For sample preparation TZ-3YB-E pellets of 11mm diameter and 3 mm height were die compacted at 200 MPa, and afterwards manually ground to 2 mm x 2 mm x 7 mm rectangular specimen. Different heating rates in the range of 0.5 K/min to 20 K/min were used up to a maximum temperature of 820 °C and a 120 min dwell time. Elongation for the different heating rates is shown in fig. 13. The specimens expand during heating depending on the heating rate and shrink during the dwell time also depending on the heating rate. Expansion increases for heating rates from 0.5 to 10 K/min, and is constant for heating rates exceeding 10 K/ min. The elongation maxima at 150 °C and 350 °C are due to the binder removal (compare to fig. 12). Debinding produces gas inside the specimen [39] which causes stresses in the weakly connected powder network. Allowing sufficient time for the gas removal
1.
DIL 802S, Bähr Thermoanalyse GmbH, 32609 Hüllhorst, Germany
59
Direct Ceramic Machining Process
avoids crack formation in larger powder compacts. The higher the heating rate the less time is available for gas removal, the stresses become greater and therefore also the risk of crack formation. Hence, low heating rates are recommended to produce reliable, crack-free blanks.
Elongation [%]
0.4
Heating Rate 0.5 K/min 1 K/min 5 K/min 10 K/min 15 K/min 20 K/min
0.3
10 15 20
5 1
0.2 150 °C 0.1
0.5 Origin Data: presintering/pre_850.opj Image: presintering/pre_850.wmf
0.0 0
200
400
600
800
Temperature [°C] Fig. 13: Dilatometry of TZ-3YB-E green bodies using different heating rates. Maximum temperature 820 °C, dwell time 120 min.
A direct correlation of differential thermal analysis and dilatometry is illustrated in
fig. 14 for TZ-3YB-E using a heating rate of 1 K/min up to 830 °C. The first exothermic peak of the DTA at 150 °C correlates to a pronounced elongation maximum. The second exothermic DTA peak correlates with a shrinkage. Further heating increases the expansion followed by the sintering shrinkage. The complex physico-chemical processes during binder removal like polymer degradation, formation of volatile and non-volatile remnants, surface interactions and mass transport were not the subject of this work.
Direct Ceramic Machining Process
60
They are investigated and reviewed by Lewis [39, 40]. The comparison of DTA with dilatometry clearly shows the link between expansion and binder removal. In order to minimize deformation and stresses during heating treatment, low heating rates are rec-
0.20
3
0.15
2 Origin Data: presintering/DTA_01.opj Image: presintering/DTA_01.wmf
0.10
1
0.05
DTA [µV]
Elongation [%]
ommended especially the larger the blanks are.
0
0.00 0
200
400
600
800
-1
Temperature [°C] Fig. 14: Thermal behaviour of TZ-3YB-E. Comparison between DTA and dilatometry. Heating rate 1 K/min.
However, economic considerations may require higher heating rates to shorten the thermal processing. Since dilatometry, DTA and TG analysis were performed either with small specimen or small amounts of powder this limits the findings. Heat treatment of larger green bodies may require more sophisticated thermal processes than we have chosen here for experimental reasons.
61
Direct Ceramic Machining Process
Further dilatometer experiments were performed for different maximum temperatures to show the shrinkage behaviour. Specimens were prepared as mentioned before. Maximum temperatures from 700 °C to 1050 °C with a dwell time of 120 min were used at a heating rate of 1 K/min. The expansion and shrinkage is shown in fig. 15. For maximum temperatures below 875 °C the blank expands, for temperatures higher than 875 °C however the blanks show a net shrinkage. This shrinkage behaviour compares with the results found by Märki [17]. The blanks experience a higher shrinkage and therefore
Shrinkage [%] Expansion [%]
a higher density with increasing maximum temperatures.
1.0 0.5 0.0 -0.5 Origin Data: presintering/pre_t.opj Image: presintering/pre_t.wmf
-1.0 -1.5 -2.0
700
800
900
1000
1100
Temperature [°C] Fig. 15: Relative length change of the TZ-3YB-E compacts after heat treatment at different maximum temperatures and 120 min dwell time (heating rate 1 K/min).
Direct Ceramic Machining Process
IV.1.7
62
Properties of the Blanks
The properties of the blanks which will be analysed in this chapter are the average density, the homogeneity, the mean porosity and the porosity distribution, the hardness, and the machinability as a function of the compaction pressure and the maximum heat treatment temperature.
Mean density
Mean density of the blanks is controlled by heat treatment temperature and compaction pressure. Pellets of 11 mm diameter and 3 mm height were compacted1 at different pressures. The samples were heat treated2 using 1 K/min heating rates up to different maximum temperatures with 120 min dwell time. Density was calculated from dimensions and mass (appendix: chapter VI.2). The density of TZ-3YB specimen as a function of compaction pressure and heat treatment temperature is shown in fig. 16. A theoretical density (TD) of 6.05 g/cm3 for TZ-3YB was used [34, 36]. For the applied compaction pressures and temperatures the relative density ranges from 41 % to 54 %TD. The relative error was estimated to be 1.5 % for the density, and 2 % for pressure, and an absolute error of ± 5 K for the temperature. Increasing the compaction pressure showed higher densities for all temperatures. All specimens showed a decreasing density up to 800 °C due to binder burn-out as due to expansion (see fig. 15). For temperatures higher than 800°C the density increases due to sintering shrinkage.
1. 2.
Graseby Specac 15011, Portmann Instruments AG, Biel-Benken, Switzerland N7H, Tony Güller AG, Hägendorf, Switzerland
63
Direct Ceramic Machining Process
56
3.4 3.2
52
Image: sinter-properties/density/density.wmf
50
3.0 Pressure in [MPa] 50 100 200 300 400 500
2.8
2.6 2.4
0
200
400
48 46
% TD
Density [g/cm3]
54
44 42 600
800
1000
40
Temperature [°C] Fig. 16: Density as a function of compaction pressure and heat treatment temperature for TZ-3YB specimen. Heating rate 1 K/min, dwell time 120 min.
Density distribution in a blank
Density distribution in a blank is a crucial property as it controls local shrinkage and therefore determines the dimensional accuracy after sintering to 100 % of the theoretical density. A review of techniques for determining density distribution is found in [11]. X-ray computer tomography was used in non-destructive measurements [9, 10, 41]. The density distribution inside a dense sintered alumina body was determined inside a dense sintered alumina body with a local resolution of 0.025 mm3 and a relative error of 0.5 %TD of alumina1 [10]. However, zirconia in presintered as well as in dense sintered state has a high coefficient of absorption compared to alumina and therefore absorbs a high amount of x-rays making CT analysis difficult. Hence, in this work density distribution was analysed by machining rings from different positions of the cylindrical blank. However, local resolution of this method is poor compared to CT. Rings of approximately 1 mm height and with different diameters, which were defined as average of outer and inner ring diameter, were machined out of a cylindrical blank of 22 mm diam1.
TDAl2O3 = 3.98 g/cm3
Direct Ceramic Machining Process
64
eter and 140 mm length using a lathe1. The blank was fabricated of TZ-3YB-E powder with a compaction pressure of 200 MPa using the dry-bag method and afterwards heat treated at 850 °C for 120 min dwell time at 1 K/min heating rate. The edges of the rings were visually checked for break-outs. The overall density of these rings was determined using geometry2 and weight3 as function of the axial and the radial coordinate. Measurement error amounts to 0.5 % of density which equals to ± 0.01 g/cm3. Rings of an outer diameter smaller than 5 mm were not analysed due to the high measurement error. A density contour plot as functions of axial and radial distance from the origin in the upper left corner is shown in fig. 17. It shows the half height of the cylindrical blank. The density distribution is assumed symmetrical and therefore for the other half of the cylinder it is obtained by mirroring the figure on the vertical axis of the right side. Density resolution of the plot is adapted to the estimated density error. The lowest density is observed at the axial coordinate zero near the hard punch of the dry bag. Density rises strongly with increasing axial distance from the punch, reaches its maximum at about 5 to 7 mm below the punch and drops slightly before reaching its maximum value again. In front of the punch up to 10 mm in axial direction the density increases in radial direction. An outer layer on the cylinders surface show a lower density than the inner bulk. Dry-bag CIP uses hard punches and compacts the powder radially. Therefore, density distribution differs from that calculated for uniaxial compaction [42] or that shown for isostatic compaction [7]. However, the density distribution found is according to our expectations: homogeneous uniform density within the middle of the blank, and decreasing density towards the hard punch in the vicinity of the ‘elephant foot’ (see fig. 6) where the diameter of the blanks gets wider and the powder has not been compacted as dense. Axial extension of density inhomogeneity corresponds well to that of the ‘elephant foot’ of 28 mm. The 5 % density drop of maximum density in direction towards the punch in dry-bag isostatic compacted and presintered blanks clearly exceeds our allowed homogeneity tolerance. However, the middle section of the blank is the most homogeneous part of the blank and have to be evaluated if bodies made of that part meet the dimensional accuracy during final sintering. 1. 2.
3.
102-VM, Schäublin SA, Bévilard, Switzerland inner diameter: Tri-o-Bor, Brown & Sharpe Tesa SA, Renens, Switzerland (error 0.002 mm), outer diameter: DigiCal, Brown & Sharpe Tesa SA, Renens, Switzerland (error 0.002 mm), height: Compac, Geneve, Switzerland (error 0.005 mm) AE 200 Delta -Range, Mettler Toledo, Greifensee, Switzerland (error 0.002 g)
65
Direct Ceramic Machining Process
Image: d_density/cylinder_01.jpg
Fig. 17: Local density distribution in a blank prepared by dry-bag CIP compaction at 200 MPa and heat treated at 850°C.
Pore size distribution of blanks
Porosity and pore size distributions of blanks and medium pore radius as function of the compaction pressure and of the heat treatment temperature were determined in the same way as for green bodies (chapter IV.1.5) by mercury intrusion porosimetry (appendix: chapter VI.2:). TZ-3YB pellets of 11 mm diameter and 3 mm height were prepared using different compaction pressures and different presintering temperatures. Heat treatment was always performed with 1 K/min heating rate and 120 min dwell time at maximum temperature with furnace controlled cooling. Pieces of about 0.5 g were cut out of the middle of these pellets for analysis. A typical pore size distribution for a compaction pressure of 300 MPa is shown in fig. 18. For all specimens presintered at 850 °C and 1000 °C the pores were between 7 nm and 100 nm. Specimens compacted with 50 MPa showed a pronounced maximum at 40 nm to 50 nm for 850 °C and 1000 °C, respectively. Specimens compacted with 300 MPa and 500 MPa showed peaks at 25 nm and 28 nm for 850 °C and 1000 °C, respectively and a small amount of pores greater than 50
Direct Ceramic Machining Process
66
nm. A relative pore size distribution of a blank compacted with 300 MPa and presintered at 850 °C is shown in fig. 19. For this pore size distribution an additional peak is observed at 7 nm pore size which may be attributed to the packing of the primary crystals (27 nm mean diameter). The 25 nm and 28 nm pores are interagglomerate pores which
Cumulative Volume [mm3/g]
developed during calcination of primary crystals.
250
green 850 °C 1000°C
200 150
100Data: sinter-properties/porosimetry/file opj/Rad_1.opj Origin Image: sinter-properties/porosimetry/file opj/Rad_1.wmf 50 0 0.01
0.1
1
10
Pore Radius [µm] Fig. 18: Pore radius distribution in TZ-3YB compacts for different temperatures. Compaction pressure 300 MPa, heating rate 1 K/min, dwell time 120 min.
67
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-2
derivation of cumulative volume (300 MPa, 850 °C)
-4
-12 1E-3
0.01
0.035 µm
-10
0.025 µm
-6 Origin Data: sinter-properties/porosimetry/file opj/p98-25.opj Image: sinter-properties/porosimetry/file opj/ -8 0.007 µm
Derivative [arbitary units / 1000 ]
0
0.1
1
Pore Radius [µm] Fig. 19: Pore sizes in the blank ranging between 7 nm and 35 nm, with a pronounced peak at 25 nm. Derivation of the cumulative volume, blanks fabricated with 300 MPa compaction pressure, 850 °C heat treatment temperature, 120 min dwell time, 1 K/min heating rate.
The pore radius distribution is shown as example for a blank fabricated using 300 MPa compaction pressure and 850 °C heat treatment temperature. The mean pore radius of 26 nm demonstrates the nanoporosity of the blanks. The changes in mean pore radius ars shown in fig. 21 for different temperatures. Lower compaction pressure produces larger pores in the blanks. The mean pore size decreases with increasing compaction pressure up to 300 MPa and then stays constant for higher compaction pressures. For heat treatments at 850 °C and at 1000 °C it remains identical independent of the applied compaction pressure. This constant mean pores size is in fairly good agreement with the overall dimensional change behaviour which was found in dilatometer measurements (see fig. 15). A linear shrinkage of 2 % will cause a negligible change in the mean pore size due to sintering. In fig. 21 the mean pore sizes of green bodies are included for comparison. The binder content in the green body significantly reduces the measured mean pore size.
68
60
100
50
80
40
60 mean pore radius 0.026 µm
30 Daten: sinter-properties/porosimetry/file opj/histo_1.opj Origin Image: sinter-properties/porosimetry/file opj/histo_1.wmf 40 20 10 0
0.01
0.02
20
0.03
0.04
0.05
0
Cumulative Counts [%]
Counts [-]
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Pore Radius [µm]
Mean Pore Radius [µm]
Fig. 20: Pore radius distribution in a blank fabricated using 300 MPa compaction pressure, 850 °C, 120 min dwell time, 1 K/min heating rate. Mean pore radius is about 26 nm. More than 90 % of the pores are in the range from 5 nm to 35 nm.
t: green t: 850 °C t: 1000°C
0.040 0.035
0.030 Origin Data: sinter-properties/porosity/file opj/mittpr.opj Image: sinter-properties/porosity/file opj/mittpr.wmf 0.025 0.020 0
100
200
300
400
500
Compaction Pressure [MPa] Fig. 21: The mean pore radii for blanks presintered at 850 °C and 1000 °C, identical heating rate of 1 K/min and dwell time of 120 min.
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A SEM micrograph of a fractured surface of a blank sintered at 850 °C for 120 min is shown in fig. 22. The fractured surface shows a homogeneous finest grained structure without any indication of persistent interstices between the powder granules (intergranular pores). Primary crystalites were 100 nm to 150 nm in diameter, they have grown to about three times the size of the crystalites in the starting powder. SEM and visual inspection extend the porosimetry analysis to a larger length scale, in both cases a homogeneous blank is found. SEM confirms the evidence that only intragranular pores
/sinter-properties/SEM-green-body/blank-microstructure.jpg
are present in the blank.
Fig. 22: Fracture surface of a blank presintered at 850 °C for 120 min. No evidence for macro pores larger than 10 µm was found. Primary crystalites are about 100 nm. (Compaction pressure was 300 MPa.)
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Porosity analysis confirms the homogeneous distribution of pores in the blank and a narrow distribution of pore sizes. For our blanks pore sizes are less than 100 nm, and therefore achieving full density is assumed during final sintering (see porosity elimination map [1]) .
Hardness
Hardness of the blanks affects the machinability and the tool wear. The hardness was investigated as a function of the compaction pressure and the heat treatment using the Vickers indenter method (appendix: chapter VI.2). Pellets with 11 mm diameter and 3 mm height were uniaxially compacted and presintered. The heating rate was set to 1 K/min and the dwell time to 120 min. The hardness results are shown in fig. 23. For reference the green body hardness values at room temperature (RT) are included. Hardness shifts to higher values as heat treatment temperature and compaction pressure increase, the increase is moderate for temperatures below 800 °C.
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Vickers Hardness [MPa]
100 Pressure [MPa] 50 80 100 200 300 60 400 500 Image: sinter-properties/hardness/hardness_f_p.wmf 40
20 0
0
200
400
600
800
1000
Temperature [°C] Fig. 23: Vickers Hardness of blanks as function of the heat treatment temperature and the compaction pressure. Relative error of hardness is estimated to smaller than 2 %.
For handling of the blanks high hardness is preferred and therefore high heat treatment temperatures should be chosen. However as hardness increases the machinability decreases and the tool wear becomes high. Therefore lower heat treatment temperatures might be preferred.
Machinability
Machinability producing exact shapes with low tool wear and high material removal rates is a crucial property for blanks. It is a very complex feature [43] and defined as the integration of all material’s properties that influence its machining process. A unique value to describe machinability has not been found up to now. Hence, machinability is expressed by four criteria: tool life, machining force, surface quality, and chip-
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ping including their size and shape. For metals this has been widely researched, for machining ceramics in presintered state only little is known. In this work the term machinability is used in a simplified way: surface quality, edge quality, and shape and size of chippings are the criteria to be analysed. Klocke et al. [44, 45] and Schippers [46] analysed machinability for green and presintered compacts by variation of tool design, tool material, cutting depth, and feed rate. Surface quality (roughness Rz) was found to be independent of the cutting depth and cutting speed. However, increasing the feed rate deteriorates the surface quality. From a tool materials viewpoint diamond, PCD or CBN exhibit lowest wear rates e.g. best machinability for machining green alumina. However, these authors approached the green machining from a production engineers’ point of view and, hence, reported no experiments varying materials processing parameters for machinability of presintered oxide ceramics. Song et al. [47] analysed the machinability of two different zirconias1 as a function of compaction pressure and presintering temperature. They found that presintering temperatures up to 1000 °C do not deteriorate edge quality, and concluded that machinability is only a direct function of the initial defect size in the blank. Furthermore both zirconias used in their analysis showed very poor edge retention. For machinability analysis concerning the surface roughness and the edge quality TZ-3YB powder was compacted using different CIP compaction pressures. One series was left in green state, the others were heat treated at different temperatures ranging from 700 to 1050 °C (50 °C increase) using a 1 K/min heating rate, and a 120 min dwell time at maximum temperature. Temperature was measured directly at the blanks with thermocouples type B. The blanks were carefully clamped in a turning machine2 in a rubber-lined three-jack chuck. The cylindrical surface was traversed using 2800 rpm and 950 rpm at feed rate of 75 mm/min and subsequent cutting depths of 1.0 mm, 0.5 mm, 0.1 mm and 0.02 mm with a hard metal tool3. The top circle surfaces of the cylinder were surfaced using a rotational speed of 950 rpm and 2800 rpm with a manual feed. A HM tool was used for surfacing. Surface and edge quality were checked visually using 1. 2. 3.
SY-Ultra 5.2 supplied from Z-Tech, Australia and YZ5N supplied from Tioxide Specialities Ltd., UK Weiler, Krucker, 8600 Dübendorf, Switzerland Sandwik: CCMT 06 02 08-UR H13A (rough turing tool), DCMT 07 02 08-UF H13A (fine turning tool), Brütsch / Rüegger AG Werkzeuge, Zurich, Switzerland
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light stereo microscopy1. The surface roughness and the edges of a machined blank are shown in fig. 24 as a function of the heat treatment temperature and the cutting speed (rpm) for a constant 300 MPa compaction. For 850 °C the cylindrical surface and the top circle surface are very smooth - roughness is hardly perceptible. The edge is also very smooth, clearly defined and no uncontrolled chippings were observed with higher magnification. In contrast, on samples presintered at 1050 °C high roughness and severe chippings of approximately 0.5 mm diameter were visible on all surfaces. Increasing the heat treatment temperature from 850 °C on to 1050 °C gradually increases the roughness on the cylindrical surface and the top circle surfaces. The edges also show gradually more chippings as heat treatment temperature is increased. Blanks heat treated at 800 °C and below as well as green bodies were also machined using the same parameter set. However, they are not included in fig. 24 because they were too weak to withstand the cutting forces and broke. Therefore, it is concluded that appropriate heat treatment temperature is at 850 °C to 900 °C. Technological data for turning machining were not optimized. However, surface roughness is equal for both rotational speeds used here. For the lower rotational speed the grooves of the tool are clearly visible on the cylindrical surface.
1.
Leica Wild M10, Leica Microsystems AG, Glattbrugg, Switzerland
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(a)
74
smooth
(b)
increasing surface roughness and edge chippings
850 °C
900 °C
950 °C
1000 °C
1050 °C machinability_03.jpg
machinability_01.jpg rough
Fig. 24: Optical micrographs showing machined surfaces and edges of blanks. Roughness and edge chipping increased within increasing presintering temperatures. Machinability as a function of the presintering temperature for constant compaction pressure of 300 MPa. (a) turned cylinder surfaces by traversing for two rpms: left side for 950 rpm, right side for 2800 rpm. (b) view on the top circle surfaces and the edges created by surfacing at 2800 rpm and manual feed.
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Machinability concerning surface roughness and edge quality was also analysed as a function of the compaction pressure for a constant heat treatment at 900 °C. The blanks were prepared as described before using different compaction pressures. The results concerning the surfaces and the edges are shown in fig. 25. For all compaction pressures smooth cylindrical surfaces and top surfaces are apparent. Pittings of small size during machining are not judged to be of importance (black arrows). The edge of the 100 MPa sample seems to be of slightly higher quality e.g. lower roughness than those with higher compaction pressures, however the difference to the 300 MPa is very small. The cutting speed had a negligible effect.
(a)
(b)
minor effects on surface roughness and chipping of the edges
100 MPa
200 MPa
300 MPa machinability_5.jpg
machinability_7.jpg
Fig. 25: Machinability as a function of the compaction pressure for blanks heat treated at 900 °C. Machining tools and parameters were kept constant except the cutting speed (rpm). (a) turned cylinder surfaces by traversing for 950 rpm and 2800 rpms. (b) view on the top surfaces and the edges created by surfacing at 2800 rpm and manual feed.
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Machinability of the blanks depends mainly on the heat treatment temperature whereas the compaction pressure has only a weak influence. These findings are in contrast to those by Song et al. [47]. In conclusion: For fabrication of blanks the compaction pressure is chosen at 300 MPa and the heat treatment temperature at 850 °C. Turning and milling are governed by the same material removal processes [43]. Therefore the results obtained for turning are also valid for milling. Machinability also depends on the shape and size of the chippings produced during the materials removal process at the places where the cutting edges of the milling tool interact with the blank. Chippings produced when using different milling tools were analysed. The blanks were compacted at 300 MPa and presintered at 850 °C for 120 min dwell time. Milling was performed using a ball-ended milling tool with 1.5 mm diameter at 18’000 rpm with a 400 mm/min continuous feed rate and a 6.0 mm cutting depth. In addition a second tool of 3.0 mm diameter was used also at 10’000 rpm with a 200 mm/ min continuous feed rate and a 6.0 mm cutting depth. Milling using the small tool corresponds to a fine or finish milling operation whereas using the larger tool corresponds to roughing. Rough milling produces a higher material removal rate. Fine milling produces accurate dimensions and smooth surfaces. Chippings were collected and for comparison also chippings produced by milling a green body. The results are shown in fig. 26. The chippings vary in size and form but only few chippings are larger than 100 µm. The majority of chippings are small fragments of a few microns. The shape of the chippings is irregular and the edges are sharp. This is typical for brittle separation of material. No evidence has been found for cutting or for plastic deformation. Therefore removal mechanism is due to chipping-off particles at their weakest linking points as expected for brittle materials. The chippings produced using the larger tool exhibit similar irregular shapes and sharp edges like the chippings produced with the small tool. However, the size of rough milled chippings is larger. Hence, both operations are controlled by the same material removal mechanism because of the similar shape and edge characteristics of the chippings.
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(b)
(c)
(d) machining/milling_4.jpg
(a)
Fig. 26: Chippings produced by milling a blank using tools of different diameters in full cutting mode. Top row: Chippings from a ball-ended milling tool with 1.5 mm diameter, 18’000 rpm, 6.0 mm cutting depth, 400 mm/min feed rate. Bottom row: Chippings from a ball-ended tool with 3.0 mm diameter, 10,000 rpm, 6.0 mm cutting depth , and 200 mm/min feed rate.
The machinability of green bodies in terms of chipping size and shape shows some interesting differences compared to machinability of blanks. Green bodies fabricated using 300 MPa compaction pressure were machined with the same tools and machining parameters as mentioned before. The chippings were analysed and the results are shown in fig. 27 for the small diameter tool, and in fig. 28 for the large diameter tool. The size
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of the chippings range from a few microns to approximately 100 µm for finishing (1.5 mm diameter tool), and from few microns to more than 200 µm for roughing (3.0 mm diameter tool). The fraction of small sized chippings in case of green milling seems to be less than for milling in presintered state. Different material mechanisms are likely to produce the chippings during green milling. Material removal due to cutting mechanism is shown in fig. 27 (b, d): the white arrow points to a surface marked by friction grooves of the tool whereas the black arrow points to a surface created by brittle fracture. That special chipping might be created by a sequence of different mechanisms: first by cutting which involves ductile deformation due to the cutting edge of the tool penetrating into the green body. Secondly, the chipping separates due to brittle failure from the green body when its limiting strength is reached and no further deformation is possible. Furthermore, fig. 27 (c) shows a chipping which was created by cutting (white arrow) and the rest of the chipping is similar to a lamellar or a Scher-chip such as observed during metal turning [43]. For the 3.0 mm diameter tool the same characteristics are observed and shown in fig. 28. Therefore, there is strong evidence that different mechanisms control material removal depending whether the machining is carried out in the presintered or in the green state.
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(a)
(b)
(c)
(d)
machining/milling_5a.jpg
Fig. 27: Chippings produced by milling a green body using a ball-ended milling tool with 1.5 mm diameter, 18’000 rpm, 6.0 mm cutting depth, 400 mm/min feed rate in full cutting mode. (a) Size of chippings ranges from about 100 µm to many fragments of a few microns. (b, d) Surface originated by cutting mechanism (white triangle) and by brittle failure (black triangle). (c) Plastic deformation of the green material.
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machining/milling_6.jpg
80
Fig. 28: Chippings produced by milling a green body using a ball-ended milling tool with 3.0 mm diameter, 10,000 rpm, 6.0 mm cutting depth, 200 mm/min feed rate in full cutting mode. Chippings range from large size of more than 200 µm to fragments of few microns. Material removal mechanisms are driven by cutting and brittle failure. Plastic deformation of the chips is observed, too.
A model for the different material removal mechanisms is presented in fig. 29. The chipping during milling of blanks are created by a brittle failure, and show irregular sized and shaped fragments. Organic binder are already burnt out of the blanks. Therefore the blank consists of the compacted and deformed network of particles which have created brittle interparticle connects during the presintering process. These interconnects show brittle fracture and are the weakest link in the chain while moving the cutting edge through the blank. Therefore, cutting is not observed for milling of blanks. In contrast, milling green bodies produces chipping surfaces with cutting grooves and chippings which show a high level of ductile deformation. Ductile deformation to some extent and cutting is only possible due to the organic binder matrix in which the soft but compacted spray granules are embedded. Strong interparticle connection is not yet established and the granules can be split off in subagglomerates. Moving through this structure the cutting edge can easily separate chippings either through the binder matrix or the spray granules (fig. 29(a)). Brittle failure is difficult to control in order to achieve the required accuracy without any damage in the machined part. Cutting is much easier to control because the ability of ductile deformation reduces the risk of unintentional damage to the part. In terms of machining, cutting mechanism is preferred and therefore milling green bodies. However, blanks withstand higher cutting forces and enable faster sinter cycles
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for the end user. Therefore we have chosen the milling of presintered blanks for our further work. Nevertheless, the possible advantages and disadvantages of using green blanks for the milling of frameworks need further evaluation. There might be future economic reasons to choose this option which will not, however, elaborated further in this work.
machining/3-1.fm
(a)
(b)
Fig. 29: Model for the different mechanisms of material removal. (a) Milling a green body , the cutting edge separates and / or deforms the soft spray dried granules along the dashed line. (b) Milling a blank the cutting edge chips off agglomerates at the weak interparticle links.
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IV.1.8
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Colouration
All-Ceramic dental restorations should match the colour of the natural teeth in order to fulfil aesthetic demands. Since the colour of TZP is white to ivory it has to be modified by colouring the blanks. A partial adjustment of colour was possible by covering the TZP framework with coloured porcelain. However, better aesthetic properties may be achieved by colouring the TZP framework itself. Thus, a method for colouration of TZP was investigated able to produce colours close to that of natural teeth. Principally one could introduce the colouration agents during the TZP powder fabrication process [48, 49], or introduce the colouration agents in the spray-granule and calcinated state [48-51]. Furthermore it is possible to introduce the colouring dopants in the green body or in the blank [52, 53]. Lemaire heat treated the samples after sintering under reducing conditions to produce metallic nanoparticles which were responsible for evolving colour. Mitsuaki impregnated either the green body or the blank with a dopant consisting of soluble metal salts [52, 53]. The colour develops during the final sintering step. All these methods either rely on the powder manufacturer or require the preparation of slurries from which pressable powders must be produced. In our work we used infiltration of blanks with dopants as aqueous solution and subsequent sintering. Thirteen different dental colours in TZP were produced using mixtures with different ratios of iron oxide Fe2O3, bismuth oxide Bi2O3, and cerium oxide CeO2 [50]. Ivory-coloured TZP was achieved when using mixtures of different ratios of yttria Y2O3, erbium oxide Er2O3, praseodym oxide Pr6O11, iron oxide Fe2O3 and zinc oxide ZnO [48, 49]. Red coloured TZP was fabricated by a dispersion of metallic copper nanoparticles in the matrix [51]. The goal of this subchapter is to evaluate the colouration of blanks using the infiltration method, to analyse colouration effects of different cations, and to investigate the influence of concentration, infiltration time, and sintering process. Since the size of the blanks might be important for infiltration, we used different samples, either small cylindrical discs or large blanks. Furthermore the sintering behaviour of the doped TZP was analysed by dilatometry. Covering the TZP with a veneering porcelain might change the colour of the composite. Hence, two different veneers were applied in various thicknesses.
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Colouration with different cations
Colours evolving in TZP were analysed as a function of the different dopants and their concentration in the solution. For that blanks were sliced in discs of 24 mm diameter and 3 mm height and stored for 180 min in aqueous solutions of metal cations. Then, the discs were dried for 12 hours at 120 °C1 and afterwards sintered. For colouration salts of Pr(NO3)3 • 5 H2O, Fe(NO3)3 • 5 H2O and VOSO4 were completely dissolved and diluted 0.008 mol-% to 0.17 mol-% in distilled water. Concentrations and pH of the aqueous solutions are listed in tab. 4. The praseodym-nitrate solution is faint green independent of its concentration, the iron-nitrate shows yellow for small concentration and a brown-orange colour for the higher concentrations and the vanadium-oxysulphate changes from faint blue to deep blue colour with increasing concentration. The discs were sintered using a 3 K/min heating rate, 1100 °C to 1500 °C sintering temperature, and 120 min dwell time. All specimen were furnace cooled. Colour was determined in the CIELAB2 colour system using the Minolta colorimeter3 (spectrophotometry in reflection). For an introduction to the CIELAB colour system see [54, 55]. Reflection spectra were also obtained using the Minolta colorimeter. Each measurement was performed always using the same white background, and calibrating always with a white standard.
Tab. 4: 1. 2. 3.
concentration [mol-%]
Pr3+ pH
Fe3+ pH
V4+ pH
0.008
6.0
2.5
2.7
0.016
5.5
2.4
2.5
0.027
5.3
2.3
2.3
0.040
5.3
2.1
2.2
pH-values of the colouring solutions for different cations and concentrations.
Heraeus Instruments, Zürich, Switzerland Commission International d’Eclairage - L*a*b*, 1976 CM-2002, Minolta, Dietikon, Switzerland: light type D65 (mean daylight), measuring type d/8 (diffuse lighting, 8° observation angle), measuring area 8 mm diameter,
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One series of discs was infiltrated with different doping cations using 0.008 mol% concentrated solution, dried and sintered afterwards at 1500 °C. The reflection spectra are shown in fig. 30. Different doping cations produced different reflection spectra e.g. colours in sintered TZP. The visible colours are yellow for praseodym-cations, yellowgreen for vanadium-cations, and brown for iron-cations.
100
Reflection [%]
80 60 coloration/Pr_Fe_V_constant_concentration.wmf 40 Praseodym Vanadium Iron
20 0 400
500
600
700
Wavelength [nm] Fig. 30: Reflection spectra for TZP infiltrated with different cations, concentration 0.008 mol-% [56].
Colouration with different concentration
Colours as function of the dopant’s concentration were studied for praseodym, iron and vanadium. Presintered discs were infiltrated, dried and afterwards sintered at 1500 °C. The results are shown in fig. 31, fig. 32 and fig. 33, respectively. Increasing the dopant concentration e.g. for praseodym lead to a parallel shift of the S-shaped curves. Lower reflection for wavelengths up to 650 nm was observed for higher concentration. However, the visible colour of the specimens was yellow. For increasing concentration
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the reflection spectra was (a) shifted to longer wavelength which means a growing tinge of red, and (b) shifted to smaller reflection values which means the specimens look darker. Similar behaviour was observed for iron (fig. 32) and vanadium (fig. 33) although the parallel shift of the curves towards lower reflection and to larger wavelength is less pronounced. Since the dopants studied produce different reflection spectra with different dependency on the concentration, adequate concentrations must be found for each dopant to match teeth colours.
100
Reflection [%]
80 coloration/Pr_different_concentrations.WMF
60 increasing Pr3+ concentration
40
0.008 mol% Pr3+ 0.016 mol% Pr3+ 0.040 mol% Pr3+ 0.170 mol% Pr3+
20 0 400
500
600
700
Wavelength [nm] Fig. 31: Reflection spectra for praseodym in TZP as function of its concentration [56].
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100
Reflection [%]
80 coloration/Fe_different_concentrations.WMF
60 40 0.008 mol% Fe3+ 0.016 mol% Fe3+ 0.027 mol% Fe3+ 0.040 mol% Fe3+ 0.170 mol% Fe3+
20 0 400
500
600
700
Wavelength [nm] Fig. 32: Reflection spectra for iron in TZP as function of its concentration [56].
100
Reflection [%]
80 60
coloration/V_different_concentrations.WMF
increasing V4+ concentration
40 0.008 mol% V4+ 0.016 mol% V4+ 0.027 mol% V4+ 0.040 mol% V4+ 0.170 mol% V4+
20 0 400
500
600
700
Wavelength [nm] Fig. 33: Reflection spectra for vanadium in TZP as function of its concentration [56].
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The colours achieved by the different dopants praseodym, iron and vanadium are presented in the CIELab system (fig. 34). L*a*b* values of the vanadium doped samples show hardly any dependence on the concentration. For the praseodym and iron doped samples L*a*b* change with the variation of their concentrations. Increasing the praseodym concentration leads to a growing tinge of red and yellow. In case of iron the tinge of blue and green increases with cation concentration. An exception is the 0.008 mol-% iron concentration which has a*- and b*-values close to those of the 0.008 mol-% praseodym doped sample. Furthermore, the L-value decreases which means that the specimen get darker as the cation concentration increases, independent of the cation type. Different dopants behave differently when increasing their concentration.
more green
more red
coloration/Lab_Pr_Fe_V_different_concentration.wmf
55
Vanadium more yellow
45 Praseodym 40
more blue
b*-Value [-]
50
0.008 mol%
35 Iron 30 25 -2
0
2
4
6
8
10
a*-Value [-] Fig. 34: Colour of sintered TZP containing different concentrations (0.008, 0.016, 0.0271, 0.040 and 0.170 mol-%, respectively) of Pr3+, Fe3+, and V4+ cations. For Pr3+ and Fe3+ the arrow-head points in direction of increasing concentrations [56].
1.
For Pr3+ this concentration was not used.
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Colouration dependence upon sintering temperature
For this investigation presintered discs were prepared as described before but the
sintering temperature was varied in the range from 1200 °C to 1500 °C. The results of colouring are shown in the CIELAB-system in fig. 35. Vanadium doped specimens exhibit a growing tinge of yellow e.g. growing b*-value at approximately constant a*-value, for increasing sintering temperature. In case of praseodym the colour changes towards more red and yellow up to 1400 °C (a*- and b*-value), for 1500 °C both values decrease slightly. Temperature increase in case of iron doped TZP has no influence on the b*-value whereas the red colour tinge slightly increases from 1200 °C to 1300 °C and decreases again for 1500 °C. Different dopants therefore exhibit different colour if sintering temperatures are increased. Control of sintering temperature is necessary to guarantee the evolving colour for the end user.
more green
more red
55
Vanadium
40
1500°C
Praseodym
1400°C
1300°C
35
more blue
b*-Value [-]
1400°C
45
1300°C
30
1400°C 1200°C
25 20
more yellow
1500°C Lab_Pr_Fe_V_sintering_temperature.wmf
50
1200°C
-5
0
1200°C
5
Iron 1300°C 1500°C
10
a*-Value [-] Fig. 35: Colour of TZP infiltrated with equalivalent concentration (0.040 mol-%) of Pr3+, Fe3+, and V4+, sintered at 1200, 1300, 1400, and 1500 °C, respectively [56].
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Sintering characteristics of doped TZP blanks
The influence of colouring dopants on the sintering characteristics of TZP was investigated using dilatometry. Cylindrical, presintered discs were prepared as described before and infiltrated with Pr3+, V4+ and Fe3+ using 0.040 mol-% concentrated aqueous solutions and 3 hours soaking time. Afterwards they were manually ground to 2 mm x 2 mm x 7 mm rectangular bars. Heating rate was 3 K/min up to a sintering temperature of 1500 °C with 120 min dwell time. The samples were cooled down in the furnace. Pure TZP was included as a reference material. The dilatometry results in terms of relative shrinkage and shrinkage rate are shown in fig. 36. Only minor changes in relative shrinkage and sinter begin were found for praseodym and iron as compared to pure TZP. In case of vanadium doping the relative shrinkage is much lower than for the other specimens. Furthermore the vanadium doped specimen does not sinter to full density. Sintering kinetics are changed by all three dopants compared to pure TZP: iron slightly lowers the start of sintering, exhibits the highest sintering rate and the lowest maximum sintering rate temperature whereas praseodym shows a smaller sintering rate with a higher maximal sintering rate temperature. Vanadium shows the slowest sintering e.g. the smallest maximal sintering rate. Therefore all dopants even in small concentrations affect the sintering behaviour of TZP. However, in case of small amounts of praseodym and iron dopants the sintering cycle can be the same as used for pure TZP whereas it has to be adjusted in case of vanadium.
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0 Shrinkage [%]
-5
TZP Vanadium Praseodym coloaration/sinter.WMF Iron
-10 -15
Shrinkage Rate [µm/min]
-20 -25 0 -5 -10 -15 -20 -25 -30 -35
Vanadium Praseodym Iron TZP
0
200
400
600
800
1000
1200
1400
Temperature [°C] Fig. 36: Shrinkage and shrinkage rate for TZP blanks with different dopants, 3 K/min heating rate, 120 min dwell time.
Phases of doped TZP
Phase changes of doped and pure TZP were investigated before and after heat treatment as well as after sintering using XRD1. Presintered blanks were infiltrated with Pr3+, V4+ and Fe3+ using 0.040 mol-% containing aqueous solutions. Samples were produced using 1500 °C sintering temperature, 3 K/min heating rate, and 120 min dwell time. XRD did not show any additonal phases except t-ZrO2 also in all coloured samples after sintering at 1500°C for 120 min.
1.
Siemens Diffractometer D5000, Siemens AG, Munich, Germany
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Infiltration of aqueous colour solutions in larger blanks
For investigating whether full and homogeneous colour penetration of blanks with colouring dopants can be achieved larger samples of 24 mm diameter and 42 mm height were prepared. The blanks were stored in various dopant solutions for different periods of time. Afterwards the specimens were cut axially. Sintering was performed at 1500 °C with 3 K/min heating rate and 120 min dwell time. The dimensions of the cylindrical samples were 18.9 mm in diameter and 33.1 mm in height after sintering. The results are shown in fig. 37 and prove that the penetration depth of dopants is different as indicated by the colour intensity in the sintered bodies. Different dopants e.g iron and praseodym achieve different penetration depths for the same concentration e.g. 0.008 mol-% and 0.04 mol-% after 3 d soaking time. Increasing dopant’s concentration from 0.008 mol-% to 0.04 mol-% leads to greater penetration depth as was found also for longer soaking times. In case of praseodym the colour penetrated homogeneously through the whole blank. For iron doping the applied parameters did not result in a full and homogeneous penetration of the colour through the blank. The effect of dopant concentration on dependence of the penetration depth was investigated. Experiments were carried out to analyse the water-pick-up as a function of the soaking time. Blanks immersed in distilled water at room temperature showed complete penetration after 2.5 hrs, whereas blanks stored in 80 % humid atmosphere at 45 °C did not show any significant water-pick-up. Therefore saturation of the blanks with water before infiltrating can definitely be ruled out. Moreover evacuation of the blank did not show any improvement in penetration. We assume that the dopant possess a lower penetration speed than the solvent which determines the coloured zone. To our knowledge kinetics of the penetration in TZP may be triggered either by diffusion or by chromatographic effects. The sharp colour transition after sintering for low dopant’s concentration and the fact that the mobility of the dopant increases with rising temperature rule out the diffusion mechanism. Rising temperature evaporates the solvant water and therefore immobilizes the cations in case of the chromatographic mechanism. There is evidence for the chromatographic effect, however, a clear explanation of the driving force of the penetration and its dependence on the dopant has not been established so far.
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Colour evolution and the penetration depth are controlled by the dopant species, its concentration, and the soaking times. Hence, adjusting the colour of the sintered framework according to dental requirements and a guarantee for homogeneous and complete pene-
penetration_depth.jpg
tration of the blank require further investigations.
dopant: concentration: soaking time:
Fe3+ 0.008 mol-% 3d
Fe3+ 0.04 mol-% 3d
Pr3+ 0.008 mol-% 3d
Pr3+ 0.04 mol-% 3d
Fig. 37: Penetration depth of different dopants for a constant storage time (three days) in aqueous solution of different concentrations of praseodym and iron.
Colour of veneered blanks
Porcelain veneers on coloured TZP frameworks usually applied and also determine the colour of dental restorations. For these investigations dense sintered cylindrical specimens of 19 mm diameter and 2 mm height were coated either with the experimental porcelain (W35/11) according to chapter IV.2.2 using three firings, or with the commercially available IPS Empress21 porcelain. IPS Empress2 was applied according to the manufacturers instruction. Dense sintered uncoated pure TZP specimen were included as reference. L*a*b* data and reflection spectra were measured using the Minolta device. The results in terms of reflection spectra are shown in fig. 38 and in the CIELab system in fig. 39. The amount of reflected light decreases with increasing porcelain (W35/11) thickness as shown in fig. 38a. Hence, the appearance of the veneered TZP sample gets 1.
Ivoclar, Schaan, Principality of Liechtenstein
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darker as the porcelain thickness grows. Colour in terms of L*a*b* shows discontinuities with increasing W35/11 thickness and a tendency to more yellow and more red. The IPS Empress2 coating also decreases the reflection of the coated specimen in comparison to pure TZP. However, L*a*b* values change continuously with increasing porcelain thickness with pronounced red and blue colour components.
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(a)
94
100 TZP
Reflection [%]
80 60
0.65 1.08 1.33 1.68 1.88 2.32
W3511_different_thickness.WMF
40
increasing W35/11 thickness
20 0 400
500
600
700
Wavelength [nm] (b)
100 TZP
Reflection [%]
80
Empress_2_different_thicknes.WMF 1.10 1.33
60 1.45 2.00
40
increasing IPS Empress-2 thickness
20 0 400
500
600
700
Wavelength [nm] Fig. 38: Reflection spectra as a function of the thickness of the experimental porcelain (W35/11, top) and of the IPS Empress2 (bottom) on the TZP substrates [57].
9.0 -1.2
8.5 8.0
-1.4
7.5 -1.6
7.0 6.5
-1.8 coloration\Lab_W3511_different_thickness.WMF
6.0 -2.0
0
0.5
1
1.5
2
b*-Value [-]
9.5
-1.0
more yellow
Direct Ceramic Machining Process
more blue
more green
a*-Value [-]
(a)
more red
95
5.5 2.5
3.5
21.0
more yellow
4.0
20.5
3.0 coloration\Lab_Empress_2_different_thicknes.WMF 20.0
more blue
more green
a*-Value [-]
more red
(b)
b*-Value [-]
W35/11 Thickness [mm]
2.5 0
0.5
1
1.5
2
19.5 2.5
IPS Empress-2 Thickness [mm] Fig. 39: L*a*b* analysis as a function of the coating thickness of the experimental porcelain (W35/11, top) and of IPS Empress2 (bottom) [57].
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Summarizing the experimental results we found that colouration of sintered TZP is in principle feasible by dopant infiltration in the blank. A natural, tooth-like colour however was difficult to produced when using praseodym, iron and vanadium as dopants. Tooth-like colour requires other dopants such as reported in previous works of Cales [50] and Yoshida et al. [48, 49]. We have shown that the evolving colour shade is controlled by the dopant itself, its concentration, sintering temperature, and furthermore by the infiltration time. Full and homogeneous penetration of larger blanks seems principally feasible. Depth is controlled by the same parameters as colour. Therefore a tooth like colour and a full and homogeneous colour penetration is difficult to achieve simultaneously. Low dopant concentrations did not introduce any second phases in the tetragonal zirconia microstructure. The sintering behaviour has to be investigated furhter as because low dopant concentrations may lead to changes in residual porosity, sinter start temperature, and shrinkage rate as was shown in case of vanadium. On the other hand with praseodym and iron as dopants it is possible to use the same sintering cycle as for pure TZP. Visible colours for the infiltrated and sintered body differ from that observed for the colour of aqueous solutions of the metal salts. In the solution the salts are dissolved, and dissociated to cations and ligands. Colour evolves by absorption of certain wavelengths [58]. During sintering the metals are fully oxidized and the cations embedded in the ceramic material. Oxidation state and position in the ceramic matrix or in the grain boundary determine the evolving colour. Colours of aqueous solutions and colour within the ceramic matrix therefore are different.
IV.1.9
Characterisation of the Blank Fabrication Process
Tolerances of the blanks are important for supplying larger quantities with the appropriate quality. Although fabrication parameters are set to constant values the blanks properties will vary. The goal of this chapter is to analyse the tolerances of the blanks fabrication process concerning the density, the associated shrinkage and enlargement factor, using a larger number of laboratory-fabricated blanks. To prepare the blanks the following parameter set was chosen: 300 MPa compaction pressure, 850 °C heat treatment, 1 K/min heating rate and 120 min dwell time. Two different Tosoh1 powders, TZ-
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3YB and TZ-3YB-E from various lots were used. The blanks were fabricated by serveral persons in multiple batches. Variation of density, mass, diameter, and height were determined. Shrinkage and enlargement factor were calculated according to chapter IV.3.3, variation was calculated using eq. 1. Point of reference was the property (xi) of the heat treated blank e.g. relative changes are calculated according eq. 2.
maximum – minimum variation = ----------------------------------------------------mean
Eq. 1
x green – x presintered relative deviation of x = ----------------------------------------------------x presintered
Eq. 2
221 blanks prepared from TZ-3YB and 74 blanks from TZ-3YB-E were analysed. The results concerning density, mass and dimensions are shown in fig. 40. Specimens with severe edge chippings and those needed for other experiments were not included, they account for the gaps in the lines in fig. 40. Statistical analysis results of density are compiled in tab. 5, and box-plotted in fig. 41. Variation of density ranges from 1.9 % to 3.0 % and exceeds an estimated 0.25 % error in measurement1. Due to binder burn-out density decreases by approximately 3 % from the green body to the blank. Average density values for TZ-3YB-E are slightly higher than for the TZ-3YB. Mass loss for both powders is of the same magnitude as the density decrease and corresponds to the ignition loss as listed by the powder manufacturer. Relative variance of the manufacturer’s ignition loss (9.3 %) however can not account for the relative variance of mass loss for the blanks (about 50 %). This high variance includes also small break-outs of the edges. Geometric dimensions e.g. diameter and height of the cylindrical blanks experience a
1. 1.
TZ-3YB: 8 different lots; TZ-3YB-E: 1 lot (see appendix for lot numbers and chemical composition) rel. error of diameter 0.08 %, of height 0.04 % and of weight 0.03 % assuming an absolute error of ±0.01 mm of the calliper
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0.16 % expansion in case of TZ-3YB-E and 0.26 % in case of TZ-3YB. This deviation is exceeding by far the measurement error1. Expansion of the blanks correlates fairly well with the dilatometer results (see fig. 15) and is attributed to gases evolving in the green body during the binder burn-out which produce an internal pressure. This phenomenon was reported by Lewis [39], however it was not related to expansion ot the green body.
Density [g/cm3]
3.25 3.20 3.15 3.10 3.05
relative change to presintered state [%]
3.00 4
presintered state green state 0
50
100
150
200
250 0
presintered state green state 20
40
60
80
3
Data: process\normal.opj Image: process\normal.wmf
2
mass diameter height
1 0 -1
0
50
100
150
200
TZ-3YB Sample No. [-]
250 0
20
mass diameter height
40
60
80
TZ-3YB-E Sample No. [-]
Fig. 40: Analysis of blanks and green bodies fabricated from TZ-3YB (221 blanks) and TZ-3YB-E (74 blanks) powder.
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TZ3-YB
TZ-3YB-E
blank density
green body density
blank density
green body density
mean
3.084g/cm3
3.186g/cm3
3.104g/cm3
3.193g/cm3
minimum
3.049g/cm3
3.154g/cm3
3.031g/cm3
3.133g/cm3
maximum
3.108g/cm3
3.238g/cm3
3.124g/cm3
3.215g/cm3
variation [%]
1.9
2.6
3.0
2.5
Tab. 5:
Density tolerances in the fabrication process of the blanks using TZ-3YB (221 blanks) and TZ-3YB-E (74 blanks) powder.
TZ-3YB blanks
TZ-3YB-E blanks
shrinkage factor
enlargement factor
shrinkage factor
enlargement factor
mean
-0.2012
1.2519
-0.1994
1.2491
minimum
-0.2042
1.2487
-0.2057
1.2464
maximum
-0.1991
1.2567
-0.1977
1.2590
variation [%]
2.5
0.6
4.0
1.0
Tab. 6:
Shrinkage and enlargement factor tolerances in the fabrication process of the blanks using TZ-3YB (221 blanks) and TZ-3YB-E (74 blanks) powder.
The statistical analysis of shrinkage and enlargement factor is compiled in tab. 6 and box-plotted in fig. 41. For both powders the shrinkage and enlargement factors, and their variations differ only slightly. The enlargement factor shows a 0.6 % tolerance for TZ-3YB and 1.0 % for TZ-3YB-E, respectively. However, the variation of the enlargement factor should to be smaller than 0.17 % in case the required dimensional accuracy of 50 µm on a 30 mm long ceramic bridge has to be achieved (0.17 %). Fabrication variation exceeds the required tolerance, and hence requires to determine the enlargement factor for each single blank. The box plot of the density, shrinkage factor and enlargement factor in fig. 41 shows that the variation of the enlargement factor is approximately
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0.4 % (TZ-3YB-E) taking the blanks between the 25th and the 75th percentile. Hence, the required tolerances are not achievable by narrowing the subset of blanks. As a consequence, blank fabrication methods must be improved - otherwise each blank has to be characterized individually.
1.260 3.12
-0.198
3.10
Enlargement Factor [-]
Origin Data: process/PRO_1.opj Image: process/pro_1.wmf Shrinkage [-]
Density [g/cm3]
-0.200 3.08
3.06
-0.204
3.04 n= 231 3.02
-0.202
n= 231
n= 74
TZ-3YB TZ-3YB-E
-0.206
n= 74
TZ-3YB TZ-3YB-E
1.258 1.256 1.254 1.252 1.250 1.248 1.246
n= 231
n= 74
TZ-3YB TZ-3YB-E
Fig. 41: Variation of density, shrinkage factor, and enlargement factor of fabricated blanks containing minimal value (triangle), mean value (square), and maximum value (circle). The box presents 5th, 25th, 50th (median), 75th and 95th percentile.
However, smaller batches of blanks that were fabricated in-line without interruption showed variations smaller than the required tolerances. For those batches, the enlargement factor was not changed and remained constant. This proves that it is possible to fabricate blanks with much smaller tolerances than 0.17 % but it requires working within a stable, if possible automated process. Especially automated compaction units and larger furnaces will reduce variations to be within the required limits.
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IV.1.10
Direct Ceramic Machining Process
Summary
The DCM process relies on prefabricated ceramic blanks of high homogeneity with isotropic shrinkage in all spatial directions. The blanks require easy machining in a soft, porous, presintered state and are afterwards being sintered to full density. Blanks fabrication adjusts the properties of the blanks, hence the goal of the work described in this chapter was to provide a route for blanks fabrication with corresponding parameters and to characterise the blanks properties. As powder TZ-3YB and TZ-3YB-E were chosen which are available in sufficient quantity from Tosoh Corporation, Japan. Both powders are spray dried granules of tetragonal zirconia polycrystals stabilized with 3 mol-% yttria and contain additions of acrylic binder for compaction. The TZ-3YB-E powder has small additions of alumina. Sintering temperature is 1500 °C for TZ-3YB and 1320 °C for TZ-3YB-E, respectively. In the sintered state both materials exhibit submicron grain size, and exhibit high strength and high toughness. Blanks fabrication comprises cold isostatic pressing of a green body, trimming the green body, heat treating which are the mandatory steps, whereas subsequent colouration is an optional step. Parameters which affect the blanks properties are the compaction pressure, the heat treatment temperature, the heating rate and the dwell time. The green body exhibits a linear increase in density and hardness with compaction pressure. It has 45 % to 58 % porosity with mean pore radii of 20 nm at 300 MPa, homogeneously distributed in the range of 7 nm to 40 nm. Pore radii are generally reduced with increasing compaction pressures up to 300 MPa. Blanks possess a hardness from 2.5 MPa to 20 MPa which is 100 times lower as compared to the dense sintered TZP. Therefore they are very fragile and delicate to handle. During heat treatment debinding starts at 120 °C and is finished at 350 °C, it is an exothermic process as shown by DTA. The faster the heating are during presintering the larger and the more abrupt is the expansion and the shrinkage. Avoiding flaws and fabrication of reliable blanks therefore requires moderate heating rates of 0.5 K/min to 1 K/ min. Net enlargement of the blanks is observed for temperatures up to 850 °C and dwell times of 120 min. For higher presintering temperature blanks show a net shrinkage. Hardness of the blanks increases with compaction pressure and with presintering tem-
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102
perature whereas their mean density drops for temperatures below 800 °C and then starts to rise. Pore sizes are in the range of 7 nm to 100 nm depending on compaction pressure. These pores have to be eliminated during final sintering [1] in order to meet the ISO regulations [2] for dental restorations. Mean pore size remains constant for presintering temperatures between 850 °C to 1000 °C but compared to the green state the mean pore radius is larger. Homogeneity of blanks on all length scales was proved by pore size distribution on the nano scale, by SEM on the meso scale and by density distribution measurement on the macro scale. Machinability of the blanks depends mainly on the heat treatment temperature. The best machinability, defined as minimal surface roughness and minimal edge chippings, was found for 850 °C presintering temperature. Compaction pressure showed only a weak influence on machinability as defined above. Milling produced chippings of irregular shape and size in a brittle way. In contrast, the preferred cutting mechanism was observed only in the milling of green bodies. Future investigations therefore should be directed towards milling of unsintered blanks. Colouration proved to be feasible by infiltration of the blank with aqueous salt solutions. However natural dental colours were not achieved with praseodym, iron and vanadium as dopants. Infiltration of larger blanks was found to be possible. However, homogeneous and complete colour penetration depends on the dopant, its concentration, the soaking time, and also the sintering temperature. Achieving dental colour and complete and homogeneous penetration requires further investigation. Dopants even in small amounts influence the sintering behaviour of the doped TZP material which necessitates further adjustments. Mechanical properties of doped TZP should be studies in greater detail before application of this ceramic material. The achieved density of the blanks showed a small variation between maximum and minimum of around 3.0 % for more than 200 blanks fabricated. Even this small variation requires individual characterization of each single blank in terms of density, shrinkage factor and enlargement factor. However, industrial automated prefabrication of blanks may reduce the variations and enables overall characterisation of blanks. In any case the processing parameter set used in this work is not an optimum, even not for the TZP powders. Open tasks are the evaluation of alternatives for example forming by wet processing instead of dry compacting, chemical instead of thermical debinding and the use of other ceramic materials.
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IV.2
Characterisation of Zirconia - Veneer Bilayer Structures
IV.2.1
Introduction
All-ceramic dental bridges and crowns basically consist of two different material layers: the framework and the coating. New framework material requires new veneer material for coating in order to get a perfect adjustment of colour to the neighbouring teeth, and in order to lower abrasion to the antagonist teeth as which would occur in case of pure TZP. The veneer material must have aesthetic appearance, mechanical strength, and a thermal expansion coefficient close to that of TZP. The composite consisting of framework and veneer needs a strong coherence of both layers to each other without any delamination during fabrication or under external load. Porcelain veneer materials in case of porcelain-fused-to-metal (PFM) restorations are mostly based on feldspathic glasses. They are mechanically weak with bend strength in the range of 80 to 200 MPa [59, 60] (see fig. 1). The veneer may be prestressed in compression by adjusting the thermal expansion coefficients (TEC) of porcelain to a value slightly lower than that of the framework material. Compressive prestressing is favourable in case of the convex shaped crowns. In this case the chewing load tries to widen the inner metallic cavity creating tensile ring stresses in the crown that partially can be compensated by the prestressing [61]. In case of bridges chewing creates compression on the occlusal side and tension on gingival side of the interdental connector. The geometrical conditions at a bridge framework are very complex, and the prestressing mechanism should not be applied as it produces additional undetermined stresses on the gingival side of the interdental connector. Therefore a close match of the TEC of the veneer and the framework material is favourable. In case of porcelain-fused-to-metal restorations, veneers with a TEC up to 0.6 x 10-6K-1 lower than that of the metallic frame-
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work material are used [62, 63]. Metallic framework materials for restorations have a TEC in the range from 13.5 x 10-6 K-1 for Au-Pd alloys to 15.7 x 10-6 K-1 for Pd-Ag alloys [64]. In contrast zirconia has a lower TEC of 11 x 10-6 K-1. Therefore, the development of new veneering porcelains with lower TEC is required. A failure definition for dental restorations is very complex and not being extensively discussed in this work. However, a criterion for working purposes is needed to evaluate the experimental veneer porcelains. All-ceramic crowns usually fail in a catastrophic brittle mode [65]. In contrast, PFM restorations show multiple cracks within the porcelain layer or porcelain chippings but no catastrophic failure. Usually, restorations with cracks in the veneer layer are left in service whereas those with sharp edged chippings are repaired. In this work, the initiation of first crack in the mechanically weak veneer porcelain was chosen as failure criterion for evaluating the appropriate TEC of the veneer porcelain. For the investigation of the bend strength and the reliability of bilayer structures, a complete failure of the bar was taken as failure criterion. The goal of this chapter is to determine the bend strength and reliability of bilayer structures made of dense sintered TZP coated with experimental feldspathic veneer porcelains of different TEC. The veneer porcelain was evaluated using three-point bend tests of bilayer structures as a function of TEC of porcelain until the first crack in the veneer occurred. Bend strength and reliability of bilayer structures were analysed for one experimental porcelain (W35/11) in three-point bend test until the complete structure failed. A mathematical model for bilayers was used to predict the load bearing capacity. The simulation of variation in porcelain thickness, and of mastication load from different directions require to test various ratios of TZP to veneer either with the porcelain in tensile mode and TZP in compression mode or vice versa.
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IV.2.2
Direct Ceramic Machining Process
Material and Methods
TZ-3YB1 powder was compacted uniaxially to disks of 110 mm diameter at 32.6 MPa to a green density of 2.52 g/cm 3 (42 %TD). Afterwards the disks were pressurelessly sintered in air to full density in an electrical furnace2 using a 2 K/min heating rate to 600°C, 120 min dwell time, followed by a 3 K/min heating rate up to 1500°C, and 120 min dwell time. Cooling was performed in the closed furnace down to 250 °C. After sintering the density was determined as 6.05 g/cm3 (100 %).
TZP bars of different dimensions were fabricated from the disks using a flat grinding machine3 [66]. First the disks were glued on a metallic plate using shellac4, then the top face was ground flat using a cutting depth of 0.01 mm, feed rate of 10 m/min, wheel rotation speed of 21.4 m/s, lateral feed of 2 mm/stroke and aqueous cooling agent5. Grooves were cut in two orthogonal directions using a separation disk6 with cutting depth of 0.15 mm/stroke, feed rate of 1 m/min, wheel rotation speed of 21.4 m/s and no lateral feed (fig. 42 (a)). Afterwards the disk was released from the steel plate by heating to 120 °C, turned, and glued with the opposite side to the steel plate. This side of the disk was then flat ground using the same tools and parameters as mentioned before (fig.
42 (b)). The TZP bars were released by heating to 120 °C. The shellac was burnt out for 180 min at 600°C. Polishing was performed on a lapping machine7 with a 1 µm diamond suspension8 for the last polishing step. Afterwards, the bars were annealed at 1450 °C for 30 min (fig. 42 (c)) to heal any surface cracks.
1. 2. 3. 4. 5. 6. 7. 8.
Tosoh-Lot Z 304149B Nabertherm HT08/17, Tony Güller AG, Hägendorf, Switzerland Chevalier FSG-818 AD, Falcon Machine Tools Co., Ltd., Taichung, Taiwan Stettler, Lyss, Switzerland Grindex SC, Blaser Swisslube AG, Hasle-Rüegsau, Switzerland Diametal D54, concentration 75, binding B2, diameter 200 mm, width 9.0 mm, Diametal AG, Biel/Bienne, Switzerland, FLM 300, A.W. Stähli AG, Läpptechnik, Pieterlen, Switzerland A.W. Stähli AG, Läpptechnik, Pieterlen, Switzerland
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For TEC adjustment trial we used two types of TZP bars with 20 mm length, 6.25 mm width and 1 mm or 3 mm thickness (each n = 25), respectively. For bend strength and reliability tests we fabricated TZP bars with 20 mm length, 4 mm width and a thickness of 2.0 mm (n = 15), 1.5 mm, 1.0 mm, and 0.5 mm (each n = 30), respectively. In addition one series of bars (n = 15) made of porcelain with 2.0 mm thickness was produced.
(b) (a) steel plate zirconia plate second side: flattening (c)
first side: flattening and grooving
TZP bars after releasing Fig. 42: Grinding of TZP test bars.
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Veneer porcelains with different TEC were fabricated from mixtures of two glasses (Fiolax Clear1 and GM/DS2) which vary in thermal expansion. The chemical composition and the TEC of Fiolax Clear, of GM/DS, as well as of the mixtures W35/9, W35/ 3, W35/11, W35/4 and W35/10 respectively are shown in tab. 7. The TEC of the mixtures was calculated by linear interpolation between Fiolax Clear’s TEC and GM / DS’s TEC. Particle size distribution3 of W35/11 powder shows a monomodal distribution with median particle size d50 of 16 µm (fig. 43).
1. 2. 3.
Fiolax Klar, Schott Glas, Mainz, Germany GrundMasse für Dentin und Schneide, intermediate product of Ivoclar, Schaan, Principality of Liechtenstein Microtrac Full Range Particle Analyzer, 0.2 to 700 µm, Honeywell (now: Nikkiso Co., Tokyo, Japan)
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Fiolax Clear
GM / DS
W35/9
W35/3
W35/11
W35/4
W35/10
Fiolax Clear [wt-%]
100
0
40
20
15
10
5
GM/DS [wt-%]
0
100
60
80
85
90
95
TEC [10-6 K-1] (100-600 °C)
4.5
12.8
9.0
10.2
10.7
10.8
12.0
9.5
11.1
11.5
11.9
12.3
TEC [10-6 K-1] (calculated) Al2O3
5.0
13.2
9.9
11.6
12
12.4
12.8
BaO
1.0
2.2
1.7
1.9
2
2.1
2.1
CaO
1.5
2.1
1.9
2.0
2
2.0
2.1
Fe2O3
0.1
0.1
0.1
0.1
0.1
0.1
K2O
9.4
5.6
7.5
8
8.5
8.9
8.2
7.7
7.9
8
8.1
8.1
0.2
0.1
0.2
0.2
0.2
0.2
65.6
69.4
67.5
67
66.5
66.1
SrO
0.1
0.1
0.1
0.1
0.1
0.1
TiO2
0.4
0.2
0.3
0.3
0.3
0.3
4.2
2.1
1.6
1.1
0.5
Na2O
7.0
P2O5 SiO2
B2O3 Tab. 7:
75.0
10.5
Chemical composition and TEC of the veneer porcelains.
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6
Kaleidagraph Plot & Data: W35-11-3-plot.qdc Image: W35-11-3.wmf 80
5
60
4 40
3
16 µm
Counts [%]
7
2 1 0 0.1
1
10
Cumulative [%]
100
8
20
100
0 1000
Particle Size [µm] Fig. 43: Monomodal particle size distribution of W35/11 powder, a median particle size d50 of 16 µm.
The TZP bars were coated with porcelain to a total thickness of 4.0 mm for TEC adjustment and 2.0 mm for bend strength (fig. 44) respectively. For this purpose they were inserted from one side in the rectangular cavity of appropriate length and width of a plastic mould. Porcelain powder was mixed with "built-up liquid" 1 to an appropriate consistency, and the slurry was filled in the cavity of the plastic mould. For a proper filling the plastic mould was vibrated for a few seconds, and excess liquid was removed using a tissue. This procedure was repeated three times until no further excess liquid was observed. After drying the bar was de-moulded and fired according to tab. 82 on a meshtray. Excess veneer material was cut off in order to achieve the desired dimensions.Cutting was performed with a diamond saw, followed by grinding with diamond discs of 20 µm grain size, and then finally manual polishing using rotating clothed discs3 with dia-
1. 2. 3.
Built-up Liquid S, IPS Classic, Ivoclar, Schaan, Principality of Liechtenstein Ivoclar P95, Ivoclar, Schaan, Principality of Liechtenstein Abramin, Strues, Birmensdorf, Switzerland
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mond grains of 15, 9, 6 µm size. For a smooth veneer surface and rounded edges the bilayer bar was glazed according to tab. 8 after grinding. Firing was always performed in three cycles: at 930 °C for 1 min, then two times at 920 °C for 2 min, followed by the glazing at 920 °C for 1 min. The TZP bars of thickness 1.5, 1.0, 0.5 mm respectively were coated with W35/11 porcelain to a total thickness of 2.0 mm. In addition, pure veneer porcelain test bars (n = 15) of 2 mm thickness were fabricated using the same procedure. The TZP test bars of thickness 1.0 mm were coated with 3 mm W35/9, W35/3, W35/11, W35/4 and W35/10 porcelain to a total thickness of 4.0 mm, and those TZP test bars of 3.0 mm thickness were coated with 1 mm porcelain also to a total thickness of 4.0 mm, respectively.
cycle
stand-by temp. [°C]
heating rate [K/min]
closing time [min.]
dwell time [min.]
firing temp. [°C]
vacuum application temp. [°C]
firing
400
60
9
2
930
500
firing
400
60
9
1
920
500
glazing (after grinding)
400
60
1
1
920
no
Tab. 8:
Firing cycles for TZP test bars coated with veneer porcelain.
111
(a)
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(b)
Image: bilayers\zusammen.jpg
(c)
(d)
(e)
(f)
(g)
(h)
(i)
(k)
Fig. 44: Fabrication of bilayered TZP-veneer-test bars (Courtesy of the University of Zurich). (a) TZP bar in the cavity of the plastic mould. (b) Pouring the veneer slurry in the mould. (c) Drying the slurry. (d) Flattening the surface. (e) Demoulded bilayer on the mesh-tray. (f) in the furnace. (g) Bilayer after firing in the furnace. (h) Bilayer showing excess veneer coating at all sides. (i) TZP bars of different thickness, intermediate state of the bilayer, and ready-to-test bilayers. (k) Bilayer with veneer in tensile mode in three-point bend test.
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All bilayered test bars were tested in three-point bending1 using a cross head speed of 0.5 mm/min and a span of 15 mm. For TEC adjustment all bilayers were tested with veneer layer face down until the first crack in the veneer occurred. For bend strength half of the bilayers were tested with the veneer in tensile mode, and the other half were tested with the veneer layer in compression mode until complete failure occurred. The bend strength σ0 was calculated using eq. 3 with the load where the first crack occurred in case of TEC and the failure load in case of bend strength, respectively.
σ
3⋅F⋅l = -------------------20 2⋅b⋅h
Eq. 3
where F is the load in N, l the span in mm, b the bar’s width in mm, and h the bar’s total height in mm. For prediction of the load bearing capacity of bilayer structures, mathematical models can be in Kašpar [67], Timoshenko [68], and Lenz [69]. These models assume a linear elastic behaviour (Hook law), a pure bending load and the bar’s cross-sections remaining planar. A bilayer structure under a single transversal load F, in case of the three-point bending test, is shown schematically in fig. 45. The top layer is in compression mode, and the bottom layer is in tensile mode. The grey shaded area indicates the normal stress distribution perpendicular to the plane of the cross-section in the middle between the supports where the maximum momentum acts. Shear forces lying in the plane of the cross-section are required for load equilibrium but they are omitted for clarity. Stress σ perpendicular to the plane shows a discontinuity exactly at the interface of both layers. The maximum tensile stress σt, max for pure bending lies either in the edge fibre of the top layer (interface) or in the edge fibre of the bottom layer depending on the ratio of elastic moduli E and the geometry. In case σt, which is responsible for the failure, resides in the edge fibre of the bottom layer, as indicated in fig. 45, it is calculated using eq. 4, eq. 5, eq. 6, and eq. 8 where Iyy is the ideal moment of inertia for the crosssection of the bilayer bar. This assumption must be confirmed by failure surface analysis. For the calculation of the load bearing capacity according to eq. 7, the modulus of 1.
Schenk-Trebel, Ratingen, Germany
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rupture (MOR) of TZP and veneer were taken as 713 MPa and 64 MPa, respectively as derived from the bend tests of the pure materials (see fig. 49). The elastic moduli E for TZP and the veneer material were taken as 200 GPa and 70 GPa, respectively. The beam width b was assumed as to 4.0 mm, the total thickness (d + h) as 2.0 mm, and the span l as 15 mm. Light microscopy was used to analyse the failure origin.
F
E1
h
E2
d e
Image: bilayers\bilayer_system.wmf Design: bilayers\equations.ppt σc
σt neutral axis
σt, max
b F/2
z
F/2
l momentum shear force
Fig. 45: Mechanical system for a bilayered bar in three-point bending under load F. The bilayer consists of two layers of different materials (index 1 and 2). Load F produces compressive stress σc and tensile stress σt in the bilayer which are indicated by the grey shaded area. The neutral axis is at a distance e from the bottom of the bilayer. The hatched area shows the momentum and the shear forces as function of the length coordinate.
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F⋅l σ t, max = n ⋅ --------------- ⋅ e 4 ⋅ I yy
Eq. 4
E n = ------2E1
Eq. 5
h h 2 + 2 hd + nd 2 e = --- ⋅ -----------------------------------2 h ( h + nd )
Eq. 6
F = σ
4 ⋅ I yy ⋅ --------------max n ⋅ e ⋅ l
bh 3 d 3h(h + d ) 2 I yy = -------- 1 + n ----3- d 2 + --------------------------12 h + nd h
IV.2.3
Eq. 7 Eq. 8
Results and Discussion
TEC adjustment
The bend strength of bilayer structures as function of the TEC difference between the veneer porcelain and TZP is shown in fig. 46. bilayer veneer porcelain - TZP structures exhibited a high difference in bend strength with increasing ∆TEC for both investigated porcelain veneer thicknesses. Bend strength of the 1 mm porcelain series decreased with increasing ∆TEC, showed a minimum at ∆TEC = -0.25 x 10-6 K-1 and increased with for positive ∆TEC. The bend strength of 3 mm porcelain layer series was much lower but increases with ∆TEC. This difference in behaviour can not be explained by using eq. 3. For ∆TEC from 0 to -0.5 x 10-6 K-1 both bilayer types showed approximately similar bend strength (see grey shaded area). All bars with 1 mm veneer showed a higher bend strength than the pure veneer material. For TEC differences exceeding 1.0 x 10-6 K-1 cracks in the porcelain veneer formed during cooling from high temperature.
115
300
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W35/3 W35/4 W35/11
W35/9
W35/10
Bend Strength of Bilayers σ0 [MPa]
250 200
Image: Veneer_Porcelain\s_null.wmf Design: Veneer_Porcelain\load_bearing.ppt Origin Data: Veneer_Porcelain\s_null.opj
150 100
MOR of pure porcelain (64 MPa)
50 0 -2.5
Porcelain 1 mm Porcelain 3 mm
-2.0
-1.5
-1.0
-0.5
0.0
0.5
1.0
1.5
∆ TEC [10-6 K-1] Fig. 46: Bend strength of bilayer TZP-porcelain structures as function of difference in TEC for two veneer porcelain thicknesses. ∆TEC = TECporcelain - TECzirconia. Error bars are set to 20 %.
Stress calculation results for the edge fibre of the porcelain due to mechanical load and TEC mismatch are shown in fig. 47. The stress was calculated using the model for bilayer structures (eq. 4 to eq. 8) (for the results see appendix: chapter VI.7, fig. 123) and the residual stress evolving due to the TEC mismatch [70](for results see appendix:
chapter VI.7, fig. 124) was superimposed. The σtotal stress in the edge fibre of the porcelain, at the load where the first crack appeared, is less divergent for both porcelain thicknesses than the bend strength (see fig. 46). For both bilayer structures failure stress is in fair good agreement with the MOR of the pure porcelain. The residual stress σTEC is of advantage to create a compressive prestressing of the porcelain in case of a thin layer and if ∆TEC < 0. This reinforces the bilayer structure. In case of a thick porcelain layer the residual stress σTEC is of tensile nature and therefore weakens the bilayer structure.
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300
116
W35/9
W35/3
W35/4 W35/11
W35/10
σtotal [MPa]
250 200
MOR of pure porcelain (64 MPa) Image: Veneer_Porcelain\s_total.wmf Design: Veneer_Porcelain\load_bearing.ppt Origin Data: Veneer_Porcelain\s_total.opj
150 100 50 0
Porcelain 1 mm Porcelain 3 mm
-2.0
-1.5
-1.0
-0.5
0.0
0.5
1.0
∆ TEC [10-6 K-1] Fig. 47: Resulting stress σtotal in the porcelain’s edge fibre of the bilayer bar caused by the load at failure (first crack) of the porcelain and the TEC mismatch during cooling. ∆TEC = TECporcelain - TECzirconia. Error bars are set to 20 %.
The bend strength for low differences in TEC is a similar for bars with 1 mm and those with 3 mm porcelain veneer (fig. 46, grey region). In case of dental restorations the thickness of the porcelain veneer layer depends on the geometric conditions in the patients mouth and the design of the framework. A restriction of the veneer thickness is unacceptable for the dental application. Therefore, a similar bend strength of the bilayer structures for different layer thicknesses is desired, more than achieving maximum bend strength for one specific thickness with a given difference in TEC. Hence, we recom-
mend to use veneer porcelains with equal or slightly lower TEC than TZP’s TEC as indicated by the grey region of fig. 46, as in case of metallic frameworks. W35/11 and W35/ 4 are in principle all suitable, we decided to take W35/11 for all further investigations.
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Bend strength and reliability
The failure of bilayers in bend strength tests was always catastrophic as shown in
fig. 48. No delamination of the bilayer was observed prior to failure. In case TZP was in tensile mode the specimen failed completely, whereas if TZP was in compression mode and the veneer was in tensile mode the veneer failed prior to the TZP. In this case, the failure originates from the bottom edge fibre (fig. 48b, white arrow) as expected from the stress analysis.
(a)
(b) Image: bilayers\failure_bar.jpg
Fig. 48: Failure of a TZP - W35/11 bilayer with the veneer in tensile mode. Top layer is TZP and bottom layer is W35/11 veneer. (a) Complete failure without delamination. (b) Fracture surface with the failure origin (white arrow) in the veneer.
The bend strength σ0 of bilayers consisting of TZP and W35/11 with varying thickness ratios but of constant overall dimensions is shown in fig. 49. Pure TZP has a bend strength of 713 MPa, the pure experimental veneer W35/11 one of 64 MPa [66]. TZP’s bend strength as determined here is lower compared to the manufacturers data [34] and to literature [27]. If the veneer porcelain is in tensile mode a significantly lower strength of the bilayers is measured than for the veneer porcelain in compression mode. Indeed for the first group the bilayers bend strength remains approximately constant close to the low value of the pure porcelain. For the other group with porcelain in compression mode the bend strength is always significantly higher and increases with increasing TZP thickness up to the bend strength of pure TZP. Standard deviation as indi-
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118
cated by the error bars is approximately 10 % for each test set. Increasing the ratio of the strong material leads to higher bend strength and load bearing capacity of the bilayer [66]. These findings correspond very well with the findings for other materials like IPS
Bend Strength σ0 [MPa]
Empress [71], In-Ceram [72], Dicor MGC [72], or Procera Alumina [73].
800 600 Image: Bilayers\strength.wmf Design: Bilayers\strength_load_bearing.ppt Origin Data: Bilayers\strength.opj
400 200 0 0
25
50
75
Percentage TZP [%]
100 TZP W35/11
Fig. 49: Bend strength σ0 of 4 mm-bilayers consisting of TZP and W35/11 veneer of varying thickness ratios. Error bars correspond to the standard deviation as statistically calculated.
The load bearing capacity prediction is made using the mathematic model for bilayer structures (eq. 4 to eq. 8) and compared to the mean load bearing capacity as measured. Calculation uses the bend strength of TZP and porcelain as measured before (see
fig. 49 at the points 100 % TZP and 0 % TZP, respectively). The results are shown in fig. 50. For porcelain in compression mode, the load bearing capacity increases, and for porcelain in tensile mode the load bearing remains constant at the pure porcelain value. The calculation of load bearing capacity compares well with the measured values, except approaching towards the boundaries where the calculated curve does not match quantita-
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tively the experimental data. When the thickness of the second layer d approaches zero, then the eq. 4 becomes similar to eq. 3 except for the ratio n. Therefore, approaching 0 % and 100 % TZP where the influence of the second layer diminishes, the calculation and the measurement differ. In these cases eq. 3 must be used. In all other cases, the calculated load bearing capacity corresponds well to the measured values. The mathematical model is based on the assumption of pure bending by moments. However three point bending induces shearing forces. Therefore, strictly speaking the mathematical model is in principle not valid. This may be also a reason for the differences to the measured data. Parameter variation may provide probable directions for further development. The calculations show for PFM bilayers that with smaller ratio n, e.g. increasing the elastic modulus E of framework material and / or decreasing the elastic modulus E of veneer material, the tensile stress in the edge fibres decreases, and therefore the load bearing capacity increases. Furthermore a growing thickness of the weak veneer porcelain leads to increasing edge fibre tensile stress and smaller load bearing capacity. To conclude from the viewpoint of a high load bearing capacity, a stiff framework using material with a high elastic modulus and / or a veneer material with low elastic modulus is desirable. Moverover we recommend to use a high framework-to-porcelain ratio.
Load Bearing Capacity [N]
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120
600 500 400 300 200
Image: Bilayers\math.wmf Design: Bilayers\strength_load_bearing.ppt Origin Data: Bilayers\math.opj
100 0 0
25
50
75
Percentage TZP [%]
100 TZP W35/11
Fig. 50: Comparison of the measured (triangles) to the calculated (lines) load bearing capacity of bilayers. Error bars correspond to the standard deviation of the measured load bearing capacity.
The Weibull plots show the bend strength distributions and failure probabilities (fig. 51). They form a broad band which extends from the left outermost regression line for pure porcelain to the right outermost regression line of pure TZP. All bilayers with different percentage of TZP are found in between these two boundaries. Bilayers with TZP in tensile mode are close to the right, high strength boundary whereas bilayers with W35/11 in tensile mode are close to the left low strength boundary. The lowest characteristic bend strength of all test sets is found for the pure porcelain, the highest for the pure TZP. The Weibull modulus m or reliability e.g. the slope of the regression line of all the test sets varies in the range form 7.8 to 20.5 (see tab. 9). No correlation of the reliability to percentage of TZP or to the mode in which TZP was tested could be found. Because bridges are bilayers their strength may be found within the same boundaries as
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long as the failure type is similar to that of bilayers. In case of bridges, the surface is unpolished and may possess surfacial flaws that reduce strength. Hence, bilayer structures as fabricated and tested here may give a first hint on the strength, the load bearing and the reliability of all-ceramic dental restoration using the same materials.
Bend Strength σ0 [MPa] 50
100
300 400 500 700 1000
Image: bilayer/weib_bar.wmf Design: bilayer/weib_biegestab.ppt Origin Data: bilayer/wei_bieg.opj
1.0
ln ln (1/(1-P))
200
0.0 -1.0 2.0 mm W35/11 0.5 mm TZP compr. 1.0 mm TZP compr. 1.5 mm TZP compr. 2.0 mm TZP 1.5 mm TZP tensile 1.0 mm TZP tensile 0.5 mm TZP tensile
-2.0 -3.0 -4.0
4.0
4.5
5.0
5.5
0.99 0.90 0.80 0.63 0.50 0.30 0.20 0.10
6.0
6.5
0.05 0.03 0.02 7.0
Failure Probability [-]
2.0
ln (σ0) Fig. 51: Weibull plot for test bars of constant overall thickness as function of the TZP ratio and of the test mode.
Direct Ceramic Machining Process
tensile layer
122
compression layer
W35/11 (100 %)
Weibull modulus [-]
characteristic bend strength [MPa]
11.44
67.5
W35/11 (75 %)
TZP (25 %)
20.46
89.7
W35/11 (50 %)
TZP (50 %)
14.07
82.2
W35/11 (25 %)
TZP (75 %)
8.93
90.4
TZP (25 %)
W35/11 (75 %)
7.80
390.2
TZP (50 %)
W35/11 (50 %)
14.87
539.4
TZP (75 %)
W35/11 (25 %)
14.28
561.1
9.94
748.2
TZP (100 %) Tab. 9:
Weibull modulus m and characteristic bend strength σ of the bilayered test bars.
IV.2.4
Summary
The bend strength of bilayers was investigated as a function of the difference between the TEC of porcelain and the TZP substrate in three-point bending tests. For high mismatch of TEC a large difference in bend strength was found for both investigated porcelain layer thicknesses. A similar bend strength is observed when the TEC of the porcelain is only slightly lower than the TEC of TZP. For this reason W35/11 experimental feldspathic porcelain was chosen for all further investigations. Edge fibre stress which was calculated by means of a mathematical model for bilayer structures with superimposing residual thermal stresses due to the TEC mismatch shows fairly good correspondence with the MOR of pure porcelain. Bend strength, load bearing capacity and reliability were determined in three-point bending tests for different TZP ratios where TZP is either in tensile or in compressive mode. MOR for the pure TZP was determined as 713 MPa, the porcelain veneer showed a MOR of 64 MPa. A mathematical model was successfully applied for prediction of the load bearing capacity of bilayers which is qualitatively in good correlation with the
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measured load bearing capacities. Calculation and experiment show increasing bend strength and load bearing capacity for TZP in tensile mode with increasing TZP ratio of the cross-section. On the other hand low, nearly constant bend strength and load bearing capacity was found for TZP in compression mode even when increasing the TZP ratio. Weibull plots of bilayer test bars show a broad band between the low strength porcelain on the left side and the high strength TZP on the right side. The Weibull-modulus
m or the reliability for all tested bars was found to be of same order of magnitude.
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IV.3
Fabrication of Multiple-Unit Dental Restorations by the Direct Ceramic Machining (DCM) Process
IV.3.1
Introduction
The fabrication of dental restoration using the Direct Ceramic Machining (DCM) process will be described in this chapter. Its steps are shown in fig. 52. A plastic model of the framework of the bridge is artistically fabricated by a dental technician using a model of the situation in the patient’s mouth. Then this plastic model is digitized and enlarged to compensate for the final sintering shrinkage. Afterwards the enlarged framework is milled out of a porous ceramic blank. During the final sintering the porous framework shrinks to its final net dimensions. Finally, the sintered framework is coated with porcelain in order to fit the patients other teeth. Dental restoration must fit at the cervical margin better than ± 25 µm [74]. The gracile cervical margin is most important as here the inner shape of a bridge joins the outer shape. Small inaccuracy leads to rejects due to non-fitting of the restorations or lacking of portions of the margin. The remainder of the surface may exhibit greater tolerances as long as the wall thickness is thicker than its required minimum. These requirements must be met on the dense sintered restorations after performing all process steps. In this chapter we report the fabrication of all-ceramic dental crowns and multiunit restorations such as bridges. The goal is to analyse and to describe each of the DCM steps (see fig. 52). The fabrication of the ceramic blanks was already described in the previous chapter IV.1. The fabrication of the plastic model by dental technicians (appendix: chapter VI.6) and the preparation of teeth by dentists are known conventional techniques and therefore, are out of scope of this work.
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DCM Process
Plastic Model Digitizing
chapter IV.3.2
Enlarging
chapter IV.3.3
Machining
chapter IV.3.4
Sintering
chapter IV.3.5
Veneering
chapter IV.3.6
Dental Restoration
Fig. 52: Steps of the DCM process to fabricate all-ceramic dental restorations.
IV.3.2
Digitizing: Capturing the Shape
Introduction
Digitizing uses plastic or wax framework models manually custom-made by a dental technician in accordance with the patient’s situation. Digitizing is the acquisition of surface data of a framework resulting in an electronic representation of single and unconnected data points. This file is merely a cloud of points consisting of x-, y-, and zcoordinates rather than an analytical description of the surface. These data points are electronically processed with a computer in order to generate information for the tool path in the milling machine. Electronic data processing offers a number of possibilities such as repairing, smoothing, and calculating analytical surface representation based on B-spline-, Bezier- or Nurbs- mathematics. Nevertheless, it involves the risk of unexpected data distortion and should therefore be limited to basic operations like enlarging and
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126
tool path generation. Therefore, the quality of the data acquisition which is defined by surface data completeness, achieved accuracy and digitizing time plays an important role. Digitizing tools, the model objects, and the digitizing parameters jointly determine the quality and have to be mutually adjusted.
Digitizing methods are either mechanical tactile or optical. The tactile digitizing strongly relies on the mechanical properties like stiffness, elastic or plastic deformation of the framework model and the tribological properties of the materials combination. The optical digitizing depends on optical properties like surface colour, reflection and absorbance of the framework model. In principal both methods can be applied but each method experiences limitations concerning the object’s geometry, slope of surfaces, abruptness of transitions, length-to-depth relations. The following chapter will describe the mechanical tactile digitizing within the DCM process focused on dental application. It includes an analysis of the shape of the models, material and methods used for digitizing, and an analysis of major influences, and characteristics determining the digitizing. The chapter concludes with a description and representation of experience gained during this work. Analysis of Framework Models
The shape and size of the models to be digitized plays an important role for digitizing. In case of dental application the DCM has to reproduce the inner part of a restoration, e.g. the framework. Frameworks exhibit always an individual shape according to the different types of restoration such as crowns and multiple-unit restoration with characteristics as curvature in multiple directions, length, number and placement of abutment tooth. Digitizing must be possible for all of these restorations and it substantially determines the quality of the DCM process outcome. The following analysis of the shape of the frameworks identifies important characteristics concerning the application. A schematic of the mesio-distal cross-section of a three-unit framework of a bridge is shown in
fig. 53 together with the terms used in the course of this chapter. A crown is a single unit restoration for one retainer or abutment tooth. In case of a three-unit framework the pontic bridges the missing-tooth-gap between the retainers. The units are joint together using interdental connectors. Multiple-unit bridges are possible e.g. four-, five- or six-unit bridges if more than one tooth is replaced.
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85-90° cavital side cervical edge
cavity crown interdental connector
pontic occlusal side
Fig. 53: Scheme of a three-unit bridge framework with definitions of the terms used in the following chapter.
The inside of an crown is hollow. With this inner cavity the restoration is cemented on the natural tooth after its preparation. The border where the natural tooth and the artificial crown meet each other is called cervical edge. The wall thickness of the crowns may reach 0.3 mm or less and the cervical edges run sharp in most cases. The crown walls are generally prepared with very steep angles up 85° to 90°. Undercuts inside the cavities and undercuts in alignment of multiple cavities are not desired but occur frequently. The crown teeth may be placed within the bridge in various numbers and different configurations according to the patient’s situation The overall size of restorations depends on the teeth replaced and its curvature. The anatomical size of different teeth are listed in tab. 24 (appendix) concerning their mesio-distal width which is approximately 10 mm and their height from the cusp to the enamel-dentin edge which is about 10 mm, too. The human jaw is curved in buccal-palatinal direction in the horizontal plane as well as in cranio-caudal direction in the sagital plane. Therefore, the bridges are also curved in these both directions and may not be regarded as a straight line of interconnected teeth.
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The cross-section area of the interdental connectors of a bridge should have minimum area from the dentists point of view. Generally frameworks are designed with a high degree of separation especially in the front teeth (anterior) region. However a minimal cross-section area of 8 to 10 mm2 is required as absolute minimum for load bearing purposes. Interdental connectors of 10 mm2 in cross-section depending on the space limitations are acceptable to dentists. The gracile designs and the complex shapes of plastic frameworks are shown in
fig. 54. They are characterized by individuality, sharp running edges and abrupt transitions at the cervical edges, very steep surfaces, and small cavities. These characteristics are a challenge for tactile as well as for optical digitizing methods.
(a)
(b)
Image: digital/digital_01.jpg
Image: digital/digital_03.jpg
Fig. 54: Plastic frameworks of a single crown, three joint crowns (a, left side) and a three-unit model (a, right side). Measuring the wall thickness of an abutment tooth (0.3 mm) (b).
The cervical edges and the inner cavities are crucial for the restorations. Frameworks have to fit to the prepared teeth without a gap at the cervical edges. A maximum absolute tolerance of 50 µm [74] is allowed at the cervical edge independent of the number of units e.g. the span length of a restoration. However, relative dimensional accuracy increases with span length. For example a framework span length of 50 mm requires a relative dimensional accuracy of 0.1 % which has to be maintained during the complete DCM process.
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The inner cavity shape of a crown must fit tilt-free to the prepared tooth. A gap of approximately 50 µm for the cement (e.g. dental glue) between the inner surface and the retainer tooth has to be allowed1. Small and deep cavities limit the application of digitizing. In case of tactile digitizing the finite diameter hinders the stylus to access the cavities or to acquire concave radii. In case of optical digitizing techniques e.g. laser triangulation or structured light principles the triangulation angle between light source and the camera limit their use.
Material and Methods
Plastic frameworks were prepared by experienced dental technicians at the University of Zurich. Light curing polymeric materials such as Targis2 or Sinfony3 were used for the framework models throughout this work. Their fabrication is described in
chapter VI.6 (appendix: fig. 122). Setup for digitizing of frameworks used a milling machine (Kern) equipped with numerical controller4 (Heidenhain) as shown in fig. 55. A tactile digitizing probe, Lemoine5 system, was mounted on the z-axis of the machine and replaced the milling spindle. The digitizing probe operates in scanning mode and continuously records data during its motion6. In this setup, the micro milling machine served as a linear axis stage for moving the framework model in x and y direction, and the digitizing probe in z direction. The numerical controller of the milling machine was bypassed by two additional computer boards for recording data and controlling the axis’ of the machine.
1. 2. 3. 4. 5. 6.
Thinnest cement layers are desired, however the available products limit the gaps down to 30 µm to 10 µm [74]. Ivoclar AG, Schaan, Principality of Liechtenstein Espe Dental AG, Seefeld, Germany Kern Micro- und Feinwerktechnik GmbH & Co KG, Murnau, Germany, micro-milling machine equipped with Heidenhain Numerical Controller TNC 151 Lemoine S.A., Paris, France Software Version 3.60
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130
Image: digital\kern_02.jpg Image: digital\digitising_02.jpg
Fig. 55: Setup for digitizing. Lemoine digitizing probe mounted on a Kern milling machine. Enlarged elliptical section shows the digitizing of a three-unit bridge framework.
The principle of tactile digitizing requires the stylus to be in direct contact with the surface of the model. While sliding over the surface the stylus is deflected by friction and / or normal force. Deflection of the stylus is continuously recorded by three inductive displacement gauges, one for each axis direction. Data points of the surface are calculated by using the deflection data of the stylus, its tip diameter, its bending parameters, and the position as measured by the machine’s axis gauges.
Digitizing the whole surface of the framework requires to turn the model once by exactly 180°. This operation was manually performed a specially designed “U”-like clamping device shown in fig. 56. The clamping device consisted of an U-like steel holder with even and parallel front- and backside. Two positioning holes each on the front- and backside serve for maintaining the position before and after turning. Two
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Direct Ceramic Machining Process
throu holes enable the fixing of the device on the machine. By means of a light curing dental polymeric cement1 the models were glued at both ends to the cylindrical rods. Prior to this the framework model was aligned within the clamping device so that it showed minimal undercuts in the direction of the stylus.
cavital side
occlusal side framework polymeric glue cylindrical rod position holes throu holes U-holder
Image: digital\clamp_02.jpg
Image: digital\clamp_01.jpg
Fig. 56: Clamping device for digitizing a bridge framework model from the cavital and the occlusal side.
For digitizing frameworks ball-ended cylindrical hardmetal styli of 1.2 mm or 1.5 mm diameter respectively were used. They were operated with 0.6 mm nominal working deflection, 1.2 mm maximum deflection, and with speeds between 200 mm/min and 900 mm/min automatically adjusted by the software. Feed was set between 0.05 mm and 0.13 mm also automatically adjusted by the software depending on the contour gradient in feed direction. The digitizing pathway and the feed are shown in fig. 57.
1.
Espe Viso Gem: Espe Dental AG, Seefeld, Germany
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digitizing: 200 - 900 mm/min
feed: 0.05 - 0.13 mm
C
E
A
max y
B
D min y
y
Image: digital\digital_04.jpg x min x
max x
Fig. 57: Digitizing pathway for a three-unit bridge framework. "A" pathway entering the cavity, "B" pathway leaving the cavity, "C" and "D" pathway for crossing the sharp cervical edge, "E" pathway over smooth pontic.
As a potential alternative to tactile digitizing linear conoscopic holography1 was also evaluated. The optical system was mounted on an experimental prototype machine (appendix: fig. 126) fabricated in our laboratory. The feed was chosen in the range of 0.05 mm and data acquisition rate was varied in the range of 100 Hz to 700 Hz. Feed speed was chosen in the range of 1.5 m/min to 2.3 m/min. Accuracy of the tactile digitizing
In the following chapter the main effects on the accuracy of the tactile digitizing process are discussed based on our results. The quality of digitized data primarily depends on the shape of the stylus and its material. Furthermore it depends on the framework which requires a non-deforming material and a suitable shape ideally with soft contour transitions. Finally a set of suitable operating parameters is necessary such as feed, speed and the nominal working deflection influence accuracy. An empirical analysis of these characteristics and their impact on the digitizing process provides a basis for adaptation to specific cases. 1.
Optimet Optical Metrology Ltd., Jerusalem 91450, Israel
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Different geometries of the styli are shown in fig. 58 . They possess different shape, tip, length which has a strong influence on the digitizing quality.
a
b c d e f Image: digital\stylos_03.jpg
g Fig. 58: Different types of styli for tactile digitizing. (a-c) cylindrical styli with ruby balls of diameter 3.0, 4.0 and 5.0 mm, (d-e) conical ruby styli with tip diameter of 1.0 and 1.5 mm, (f-g) cylindrical ball-ended hardmetal styli with diameter 1.5 and 1.2 mm.
Stylus shapes like the conical ruby styli (fig. 58 d-e) are especially suited for parts which are flat and exhibit a small z length (depth) of less than 1 mm. Due to the diameter increase the conical shafts possess high stiffness against friction-force-induced bending superior to that with long cylindrical shafts (f, g). If contours gets deeper conical styli are disadvantageous because the digitizing system is not able to distinguish between deflection caused by its conical styli or by the model surface. Ball-ended cylindrical styli (fig. 58 f, g) or cylindrical styli with ruby balls (fig. 58 a-c) are of advantage for deep parts having a z length of more than 1 mm with steep walls. A horizontal to vertical surface transition with a 0.5 mm x 45 ° chamfer was digitized using different styli: ball, ball-ended cylindrical and conical ones (fig. 58: a, d, g). One digitizing line of each scan was selected to serve as example and is shown in fig. 59. Different styli produced different digitizing results for the same transition. Especially the conical stylus created a digitizing line with slopes according to the stylus geometry (fig. 58 d). Best reproduction of the transition shape was obtained for the ball-ended cylindrical stylus and for the stylus with a ruby ball. In case of cavities the tactile digitizing is limited by the finite stylus radius and an inner radius smaller than the tip radius can not be recorded.
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0 Setup:
Z [mm]
-1
conical stylus stylus with sapphire ball ball-ended cylindrical stylus
-2 -3 -4 -5
Image: digital/kante/surface/digitized_edge.wmf Design: digital\kante\surface\digitized_edge.ppt -6
-5
-4
-3
-2
-1
0
X [mm] Fig. 59: Digitizing a horizontal to vertical transition with a 0.5 mm x 45 ° chamfer. The transition serves as a simplified model for the margin of the tooth cavity. The digitizing results of the different styli are shown for different styli operated at 100 mm/min feed.
The tip diameter choice depends on the required accuracy, the model shape, and the model material hardness. It may appear evident that accuracy will be higher the smaller the tip diameter, however Hertzian pressure at the contact point increases as the tip diameter decreases. Therefore the tip causes more deformation of the model surface which in fact decreases accuracy (example of a metallic surface is shown in fig. 60). In case of plastic models deformation can hardly be detected due to its elastic reversibility. Furthermore due to the finite tip diameter limits the ability to digitize small cavities. In general the recorded surface data do not represent the real surface, this problem is known from tactile profilometry.
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Image: digital\medallie_digital_03a.jpg Image: digital\medallie_digital_02a.jpg
Fig. 60: Plastic deformation with grooves during the digitizing of a metal surface. Conical stylus (fig. 58 d), speed 100 mm/min.
Stylus bending caused by normal force and friction force also distorts the recorded data. Bending parameters were measured in each axis direction prior to digitizing, performed by moving the stylus tip towards a stiff surface. Bending parameters were automatically adjusted by the Lemoine digitizing software to compensate for the stylus bending. Friction was not analysed because dental technicians use various materials for fabricating framework models.
Dental waxes like Plastodent artline1 showed abrasion particles sticking to the stylus which deteriorates the measurement. In some cases when using wax models the thin edges at the boundary of the cavity were destroyed due to the forces created by the stylus deflection. On the other hand, mechanical stronger light curing polymeric materials such as Targis2 or Simphony3 worked satisfactory.
Contour transitions are very abrupt at the cervical edges and may lead to accuracy problems due to over-swinging of the stylus. When crossing the thin cervical edge (see
fig. 57 C, D) the stylus loses contact to the surface. In the following the stylus overswings in the direction of the movement, and due to the deflection it back-swings until it comes in contact with the surface again. The higher the speed the wider the stylus over1.
2. 3.
We used the Plastodent artline cervical, modelling undercoating, and the milling wax from Degussa-Dental GmbH, Hanau-Wolfgang, Germany, respectively as well as and the cervical, occlusal and the crown wax from Bego GmbH & Co., Bremen, Germany. Ivoclar AG, Schaan, Principality of Liechtenstein Espe Dental AG, Seefeld, Germany
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swings. If the over-swing exceeds the working deflection the digitizing process either stops with an error signal or the complete line is repeated. Speed reduction reduces the over-swing effect. Loss of accuracy due to over-swing can be minimized by alternation of the digitizing direction as shown in fig. 57. This changes the side of distortion for subsequent digitizing lines and therefore eliminates distortion caused by over-swinging, even at high speeds. For accurate digitizing results, the feed must be adapted to the surface gradient in feed direction. If the surface is steep (see fig. 57 A, B) it is recommended to reduce feed distance. If the surface is flat (see fig. 57 E), a higher feed distance is preferable for acceptable results.
Speed is crucial for digitizing time but may cause unexpected inaccuracy at sharp edges. The result of high speed crossing of a sharp transition is shown in fig. 61 and in comparison to low speed crossing (fig. 59). High digitizing speed resulted in distortion of the edge due to over-swinging of the stylus. If digitizing is conducted alternating (in negative as well as in positive y direction) the over-swinging is observed on both sides of the edge. However, machining will create an edge of good quality because the overswing effect is recovered. The Lemoine system uses automatic speed control1 by evaluating the contour of the preceding digitizing line. Sharp transitions are “foreseen” and speed is automatically reduced when approaching the transition.
1.
Renishaw’s SP600 also uses speed control.
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0 distortion
Z [mm]
-1
0.07 mm -2 -3
Image: digital\kante\edge_distortion_due_speed.wmf Design: digital\kante\surface\digitized_edge.ppt
-4 -4
-3
-2
-1
0
X [mm] Fig. 61: Distortion at abrupt transitions caused by too high speed. A ball-ended cylindrical stylus with a speed of 1000 mm/min at a horizontal to vertical transition with a 0.5 mm x 45° chamfer transition demonstrates a 70 µm distortion of the result due to over-swing effect.
Nominal working deflection was chosen at 0.6 mm which corresponds to 1.08 N force (110 g) for the Lemoine digitizing device, maximum deflection was set to 1.2 mm which produced a force of 2.15 N (220 g). Force due to working deflection of the stylus may lead to elastic or plastic deformation (see fig. 60). However, reduction of working deflection is equal to reducing reverse force which is necessary to guarantee for permanent stylus-to-surface contact. As a rule of thumb increasing the reverse force permits a higher digitizing speed. Results and Discussion
This chapter describes the results of digitizing and compare the tactile and the optical techniques. However, in this work the tactile digitizing was used predominantly. More than 50 different framework models were digitized using different parameter settings. Linear conoscopic holography, as a representative of the optical digitizing method [75-78], was used for evaluation purposes and comparison.
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Optical digitizing methods present potential alternatives to tactile methods. They are based on triangulation, interferometry, or time-of-flight measurements as described in detail by [79]. An advantage of optical methods especially for the dental application is their contact-free operation mode. Therefore, some disadvantages of tactile systems are avoided and frameworks are load-free, no deformation of the surface occurs which may distort the measurement. However, most of these optical systems were ruled out in our work because they are not capable to scan deep cavities with a small diameter-to-depth ratio. In case of conoscopic holography the optical beam diameter is much smaller than any stylus diameter possible for tactile systems, and therefore smaller inner cavities may be scanned. Optical methods operate inertia-free in contrast to tactile ones, their digitizing speed is higher. Digitizing always produces unconnected digital data points of the model surface, so-called point clouds. Important differences are found in point clouds produced by tactile and optical methods. Data points of the occlusal side of a three-unit framework model using tactile and conoscopic digitizing are shown for comparison in fig. 62. In case of tactile digitizing, feed distance decreased as the gradient of the contour in feed-direction is increased. For the example shown here, the feed varied between a minimum of 0.05 mm and a maximum of 0.13 mm per line. Furthermore as the stylus slides over the model its deflection changes. Data points are the sum of machine position and deflection, and therefore digitizing lines are not straight in each direction. In case of the optical system, the feed was constant at 0.05 mm per line and every 0.04 mm a data point was recorded. x- and y-coordinate of the data points lie on a grid defined by a starting point x0, y0 plus constant increments δx or δy. Hence the optical system produces a regular gridded point cloud independent of the analysed shape. In contrast, tactile digitizing produces irregular point clouds depending on the analysed shape: high data density is recorded where necessary, for example at sharp, steep transitions, whereas only low data densities are needed for flat surfaces. Artefacts are data points that do not belong to the surface and are a known problem in point clouds produced by optical methods. They need to be eliminated completely and reliably from the point cloud. Artefacts are not known in tactile digitizing.
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The average data density of point clouds produced by tactile and by optical system is different (see fig. 62). In case of tactile digitizing an average three-unit framework1 needs approximately 300 lines for each side (occlusal and cavital), and its point clouds consists of 20’000 to 30’000 data points depending on its shape. The average distance between the lines is calculated to be 86 µm, between the data points it is in the range from 0.14 to 0.21 mm. The average data density of tactile digitizing is about 180 dpi which is low compared to a 600 dpi laser printer. However data density of tactile digitizing strongly depends on the framework shape and a high accuracy could be achieved using a low average data density.In comparison optical digitizing using conoscopic holography (Optimet) works with a constant data density of up to 800 dpi. It is not affected by shape and therefore has to work with a higher amount of data points in order to achieve similar accuracy as tactile digitizing.
(b)
(a)
Image: digital\pcloud_4.jpg
Image: digital\bridge_01.jpg
(c)
Image: digital\pcloud_01.jpg
Fig. 62: Point clouds of the cavital side of a framework (b) as recorded by a tactile (a) and an optical (c) system.
The digitizing speed varies between 200 and 900 mm/min for the tactile digitizing system depending on the complexity of the framework and on the working deflection. Hence, the time necessary for digitizing is between 5 and 21 min per framework side. Optical digitizing works with a much higher speed of 1800 to 2300 mm/min. Total time for one framework side is less than 10 min and is dependent only on its size, the number of scanning lines and the speed.
1.
It exhibits an average length of 26 mm and an average width of 14 mm.
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The materials for fabricating the framework models present a challenge for tactile methods because they have different mechanical properties, and for optical methods as their optical properties vary. Plastic materials are mechanically stronger, show minimal debris and minimal contact deformation compared to dental waxes, but are more difficult to work with. Spraying a thin coating on the model1 may achieve an optically homogeneous surface. However, the accuracy may suffer from the unknown coating thickness. Dental technicians prefer to work with wax models which is not possible in combination with tactile systems, but with the optical systems. The accuracy of a digitizing system can be determined as static accuracy without movement or as dynamic accuracy in movement. In case of a tactile digitizing system the static accuracy was determined to ± 3 µ m by moving in air at a slow speed of 10 mm/min ten times towards a vertical and a horizontal surface. Dynamic accuracy depends on many parameters as discussed earlier and therefore is only valid for a specific case. In our case due to variable speed, feed and individual model shape the dynamic accuracy was not determined. In case of optical systems static and dynamic accuracy show minor dependence on the part itself and therefore can be taken from the manufacturer’s technical specifications to be smaller than 10 µm 2. Accuracy of both systems are sufficient for its application in dentistry. We tested various digitizing strategies (example fig. 57) especially for the cavital side. Performing two subsequent runs with feed in negative and positive x-direction, or first in positive x-direction and then in positive y-direction did not improve significantly data quality increase nor reduce digitizing process time. The same was found when the digitizing direction was with 10, 20 or 45° angle towards the machine’s x-direction.
In conclusion, tactile digitizing as well as linear conoscopy worked satisfactory for frameworks. Tactile digitizing proved superior in accuracy with low average data density over conoscopy. Its major drawback is the stylus-to-model contact because the soft dental waxes were deformed and abraded. Advantages of tactile systems is their sensitivity to the model shape, and that the data points are free from artefacts. Sharp edges cause no problems for tactile systems as long as deformation of the thin edges does not 1. 2.
For example: Diffu-Therm, developer for non-destructive-testing by dye penetrant method by Schneider Röntgentechnik GmbH, Dortmund, Germany Internet: www.optimet.com
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occur and the speed is below 1 m/min. Optical digitizing allows of high speed non-contact, deformation-free measurement. Regularly gridded point clouds are much easier to handle as they are represented by a two dimensional array. On the other hand they exhibit a homogeneous data density and hence ignore the model’s shape.
IV.3.3
Enlargement of the Digitized Shape: Compensation for Final Sintering Shrinkage
Introduction
During sintering the porous ceramic framework shrinks to full density. Therefore the sintered object is much smaller than the porous sample. In order to achieve identical dimensions as for the original-sized framework model, compensation of the sintering shrinkage is required and the object has to be machined in an up-scaled size. This chapter deals with the final shrinkage prediction based on the blank’s properties. In the following we transformed the digitized point cloud to an enlarged one based on the predicted shrinkage. Tool path calculation from the enlarged geometry and its transfer to the milling machine was performed using standard software packages that are commercially available for these tasks 1. The shrinkage process depends on the material itself, on the compaction density and density distribution, and on the parameters of sintering process [7]. A mathematical model for the prediction of the anisotropic shrinkage for each of the three sintering stage enables the compensation of shrinkage, distortion and warpage for each infinite small voxels of the blank [80, 81]. However this type of approach is not applied here. In contrast, a set of preconditions is introduced for which a simple prediction of shrinkage and enlargement is possible.
1.
Generate tool path information: Pointmaster V2.5 (Knotenpunkt, Balingen, Germany), Tracecut V22 (Renishaw, New Mills, Wotton-under-Edge, UK) Transfer data to NC-controller: TNC, Heidenhain, Murnau-Westried, Germany
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The goal of this chapter is to describe a simple model for shrinkage prediction, to derive the enlargement factor and to describe the enlargement of point clouds. The variation of the enlargement factor which was possible to achieve with our laboratory equipment is analysed statistically.
Methods
Homogeneous porous ceramic blanks of TZP were used for the DCM process. These were fabricated according to chapter IV.1 with the aim of achieving homogeneous isotropic shrinkage in each direction. The frameworks machined from these blanks need to sinter to full density which is estimated to be equal to theoretical density for TZP (6.05 g/cm3). Mean density in the presintered yet porous state is determined for each blank. Furthermore, the shrinkage is assumed to be independent of shape, e.g. thin and thick parts of the object behave similarly. Absolute shrinkage ∆l is defined as the difference between length of the presintered sample lp to that of the sintered sample ls. Relative shrinkage s can be calculated according to eq. 9 related to the length in sintered state.
ls – lp lp ∆l s = ----- = -------------- = 1 – ---- = 1–f ls ls ls
Eq. 9
The enlargement factor f is the ratio between the known length in the presintered state lp and the yet unknown length in the sintered state ls. If shrinkage is homogeneous one constant enlargement factor f is exists for all spatial directions in a blank and for all coordinates. Mass constancy during transition from presintered blank to sintered body can be assumed, therefore the enlargement factor f depends only on the mean density of the presintered blank ρp and on the density of the sintered body ρs according to eq. 10. The sintered density is not directly accessible before sintering and therefore must be assumed of being close to the theoretical density (TD in %). Statistical methods refine this prediction and TD may be replaced by the average density value of a statistically relevant amount of sintered bodies.
143
l f = ----p- = ls
3
lp 3 --- l - = s
3
V ------p- ⋅ m - = V s --m
3
ρs ------- = ρp
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3
TD
–1
⋅ 0.01
Eq. 10
The following eq. 11 allows to calculate the length in presintered state lp necessary to achieve a certain sintered length ls.
ρs lp = l s ⋅ f = l s ⋅ 3 ------ρp
Eq. 11
Three-dimensional bodies consist of a point cloud expressed by a (m x 3)-matrix (A) with m data points Xn = [xn, yn, zn]. A linear afine transformation like eq. 12 calculates the enlarged point cloud expressed by a (m x 3)-matrix (B).
B = A⋅f
Eq. 12
The equations enable the calculation of the enlargement factor required for compensation of final shrinkage for a specific blank. The point clouds of a digitized framework shape are blown up using that enlargement factor and afterwards the tool paths are generated on the increased shape. The mean enlargement / shrinkage factors and their tolerances determine whether for each blank the factors must be calculated separately or if the mean factor is sufficient for all blanks. Statistics were applied for 221 TZ-3YB blanks and 74 TZ-3YB-E blanks, respectively.
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Results and Discussion
The enlargement factor depends exclusively on measurable macroscopic values: mean density of the presintered blank and mean density of the sintered framework. The enlargement factor is independent of the thermodynamic path of sintering, and neither heating rate, nor dwell time, or cooling cycle are considered in the equations. Therefore it is a characteristic property of the blank. Prerequisite for using eq. 12, which is an linear afine enlargement of the point cloud, is isotropic shrinkage of the blank which relies on the homogeneous density distribution in the blank. The enlargement factor will be a function of the blank’s coordinates if the density varies within the blank. Then in principle eq. 12 is further valid if f is substituted by a (3 x 3)-matrix using matrix elements fij as functions of the x-, y-, z-coordinates. However, calculation of the enlargement factor is then much more difficult due to the warpage of the object. General models of anisotropic shrinkage [81] taking into account the density distribution within the blank may predict the shrinkage as a function of the location within the blank. The enlargement of the framework depends on its positioning inside the blank. The isotropic shrinkage and therefore the homogeneous density in the blank become a prerequisite for DCM. A priori the achievable end density in sintered state relies on a prediction, and is not known until the body is sintered. As an initial assumption we used the theoretical density of TZP. Statistical methods refine this assumption using a growing number of samples. However, this method strongly relies on a reproducible sintering result, which is difficult to perform at multiple locations.
Fabrication tolerances were recorded for shrinkage and enlargement factor during this work. The minimum, maximum, mean and variation are listed tab. 10. The enlargement factor is within a 0.6 % tolerance for TZ-3YB and 1.0 % for TZ-3YB-E respectively. For both materials the shrinkage and enlargement factor, and their variation differ only slightly from each other. Therefore we conclude that both materials exhibit similar processing behaviour. A small variation e.g. small tolerance of the enlargement factor is desired for industrial production eliminating individual determination for each blank. The variation as determined in this work, however, requires a characterisation per blank.
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TZ-3YB
TZ-3YB-E
shrinkage factor
enlargement factor
shrinkage factor
enlargement factor
mean
-0.2012
1.2519
-0.1994
1.2491
minimum
-0.2042
1.2487
-0.2057
1.2464
maximum
-0.1991
1.2567
-0.1977
1.2590
variation [%]
2.5
0.6
4.0
1.0
Tab. 10: Shrinkage and enlargement factor tolerances in the fabrication process of the blanks using TZ-3YB (221 blanks) and TZ-3YB-E (74 blanks) powder.
The relative dimensional tolerance of the sintered framework should to be less than 0.17 % (see chapter II.2). Assuming accurate milling of the enlarged framework, e.g. when the relative error of presintered length can be set to zero, the upper limit of the
relative tolerance of the enlargement factor also has to be less than 0.17 % (see appendix, eq. 23 and tab. 20). Fabrication tolerances for both powders (see tab. 10) exceed the relative error of enlargement factor. Therefore individual determination of enlargement factor per blank is necessary. Due to the measurement errors (appendix: chapter VI.3) the relative error for the enlargement factor amounts to 0.05 %. Hence, it can be negligible because it is much smaller than the upper tolerance limit for the enlargement factor. Density resolution of ± 0.01 g/cm3 (see chapter IV.1.7 and fig. 17) accounts to a relative error of 0.67 % in presintered density1. On this basis, assuming that the sintered density can be measured with no error, the relative enlargement factor tolerance is 0.22 % (see appendix, eq. 20 and tab. 20). This is about the same magnitude for the relative tolerance of the enlargement factor as derived from eq. 23 (appendix: chapter VI.3). Hence we conclude in retrospect that the density distribution in the blanks is sufficient for achieving the required dimensional accuracy.
1.
for 3.0 g/cm3 mean density.
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IV.3.4
146
Machining of the Ceramic Blank
Introduction
Machining is the third step in the DCM process, following digitizing of the framework shape and its enlargement. The enlarged framework is machined out of a porous blank. Machining requires high accuracy because it is the last shaping step in the DCM process. In addition to accuracy, fast machining and therefore short machining time are desirable. Machining requires the specification of the number of machining steps including the number and the type of tools, the number of machining and clamping positions and the movement of the tools in relation to the object as well as the cutting parameters. This chapter presents the machinery and the tools for cutting frameworks out of TZP blanks. A set of appropriate parameters such as feed and speed is provided.
Materials and Methods
The milling setup used the same machine1 as for digitizing but a milling spindle2 was inserted instead of the digitizing probe (fig. 55). The numerical controller was bypassed by the Lemoine3 system which controls the movements of the axes of the machine.
1. 2. 3.
Kern Micro- und Feinwerktechnik GmbH & Co KG, Murnau, Germany, micro milling machine equipped with Heidenhain Numerical Controller TNC 151 mRV 8/5, Fischer E. AG, SFJ Präzisions-Spindelbau, Herzogenbuchsee, Switzerland Lemoine S.A., Paris, France
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Image: digital\kern_02.jpg Image: machining\mach-13_a.jpg
Fig. 63: System setup for milling the TZP blanks: Kern milling machine with mounted milling spindle. Fine milling of the occlusal surface of a three-unit framework (ellipse).
Machining the whole surface of the framework requires to turn the model once by exactly 180°. This operation was manually performed by removing and turning a specially designed “U”-shaped clamping device shown in fig. 64. Machining used the same clamping device and the same principles as for digitizing. The blank was mounted stress-free using a fast glue aligned in the centre axis direction between the flat ends of the cylindrical rods. The U-holder must be stiff that no elastic energy in axial direction can be stored which would lead to damage of the framework when released.
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U-holder
(a)
Image: machining\blank_01.jpg
throu holes positioning holes cylindrical rod blank fast glue
(b)
Image: machining\clamped_02.jpg
Fig. 64: Clamping device for machining a blank. (a) Principle: a blank is glued in between two cylindrical rods. (b) Composition of the clamping device as used.
Machining of the complete surface of a dental framework required four NC programs: a roughing and a finishing program for the cavital as well as for the occlusal surface.The NC programs were generated for the enlarged framework point cloud. The working step sequence was chosen as follows: roughing occlusal, turning the clamping device by 180°, roughing cavital, changing the milling tool, finishing cavital, turning the clamping device by 180°, finishing occlusal, removing the object and manual work-over at the mounting points. The NC programs for roughing consisted of approximately 40 travels, and of 260 to 350 travels for finishing a three-unit framework of approximately 30 mm length. The travels for a three-unit bridge framework are demonstrated in fig. 65.
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feed rate 0.1 - 2.0 m/min
feed: 0.05 - 0.13 mm
Image: machining\mach_bridge_40.jpg
y x
Fig. 65: Tool travel pathway for machining of a three-unit framework. At both sides (mesial and distal) the holding points are visible.
In this work we used milling tools with different geometries and materials from various suppliers as shown in fig. 67. The milling tools consisted of hardmetal (HM) and one of polycrystaline diamond (PCD). Roughing was performed with milling tools of 2.5 mm or 3.0 mm diameter1 having the following properties: centre cutting, two cutting edges, 18 mm cutting length, 50 mm total length and 4.0 mm shaft diameter. Rotation was set to 10’000 rpm which results in cutting speeds of 1.5 m/s. Feed rate was about 0.1 m/min for the first cut and was increased to 1.2 m/min after the first travel which is the maximum travel rate of our Kern machine. A feed rate up to 2.5 m/min was also applied when using an experimental prototype machine (appendix: fig. 126). The cutting depth was set to about 10 mm and the feed was chosen as 0.8 mm. Roughing was performed alternating in conventional as well as in climb mode. For finishing we used HM tools1 with a 1.5 mm cutting diameter, two cutting edges, and 12, 14, and 16 mm cutting length. Their total length was 50 mm, and their shaft diameter 3.0 mm. Rotation was set to 18’000 rpm corresponding to 1.5 m/s cutting speed. Feed rate was set similar to the roughing step: 0.1 m/min for the first cut and increased to 1.2 m/min after the first travel 1.
Marc Sandoz S.A., Rue Stavay-Mollodin 25, 2300 La Chaux-de-Fonds, Switzerland
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for the Kern machine and to 2.5 m/min for the prototype. The cutting depth was set to perform fine milling within one level. The feed was varied in the range of 0.05 mm to 0.1 mm depending on the contour gradient in feed direction for the cavital side, and was 0.15 mm for the occlusal side. Finishing is also performed back and forth using conventional and climb milling.
a b c d
e
f
g
h
i
k
finish mills
rough mills
Image: machining\tools_01.jpg
Image: machining\tools_02.jpg
Fig. 66: Different milling tools for finishing and roughing (refer to tab. 11).
151
Tool material No. (fig. 66)
cutting diameter [mm]
cutting length [mm]
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number of cutting angle helix angle cutting edges / [°] [°] centre cutting
a
HM
0.5
4.0
30
2 / yes
9
b
HM
1.0
4.0
30
2 / yes
9
c
HM
1.5
10.0
30
2 / yes
9
d
HM
1.5
12.0
30
2 / yes
9
e
HM
1.5
17.0
30
2 / yes
9
f
HM
1.5
17.0
0
1 / nos
0
g
HM
3.0
8.0
30
2 / yes
9
h
HM
3.0
18.0
30
2 / yes
9
i
PCD
3.0
4.0
0
1 / no
0
k
HM
2.5
18.0
30
2 / yes
9
Tab. 11: Specification of the milling tools. HM tools supplied by Sandoz1, PCD by W. Metzger2. 1. Marc Sandoz S.A., La Chaux-de-Fonds, Switzerland 2. W. Metzger & Co AG, Hardmetal and Diamond Tools, Lotzwil, Switzerland
After the milling process, the bridge framework was separated from the remaining material at the right and on the left side by manually sawing or grinding with a disc tool. The holding points were worked over by manual grinding.
Results and Discussion
The milling process combines rough and fine milling of a porous ceramic framework with approximately 1.25 times enlargement compared to the original framework model. A comparison of both samples is shown in fig. 67.
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inflated framework Image: machining\bridge_dye_07.jpg plastic model
Fig. 67: Enlarged porous ceramic framework (top) in comparison to the original model (bottom). Enlargement is approximately 1.25 times.
The substeps for machining are presented in fig. 68 for roughing and in fig. 69 for the finishing. Best results were achieved with the available equipment as follows: First the occlusal side is rough milled, back and forth travelling, in three levels with 6 mm cutting depth and 0.4 mm allowance, the tool had 3.0 mm cutting diameter (fig. 66 g), a constant feed of 0.7 mm was used, and a feed rate of 800 mm/min, at 10,000 rpm. Afterwards the clamping device is turned by 180° and the cavital side is rough milled alternating in three levels. After changing to the fine milling tool with 1.5 mm cutting diameter (fig. 66 d) a feed rate of 1’100 mm/min was used at 18’000 rpm. The cavital side is fine milled in one level to the final contour with no allowance being applied. Fine milling uses variable feed ranging from 0.05 mm to 0.13 mm depending on the contour gradient. After turning the clamping device by 180° again fine milling of the occlusal side is performed in one step to the final contour applying a constant feed of 0.15 mm. Afterwards these holding were separated manually using sawing devices, and the holdings regions are manually worked over.
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(a)
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Image: machining/milling_1.jpg
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(g) Fig. 68: Rough milling process of a three-unit framework using a milling tool (fig. 66 g), 3.0 mm cutting diameter, 10’000 rpm, 0.7 mm feed, 800 mm/min feed rate. (a) Start of the rough milling. (b) Rough milling of the occlusal side, second level of three. (c) Occlusal side completely rough milled. (d) Rough milled occlusal side (view point changed to backside). (e) Rough milling of the cavital side, second level of three. (f) Rough milled framework within the clamping device. (g) Occlusal side of the rough milled framework. (Scale bars: 10 mm)
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Fig. 69: Fine milling process of a three-unit framework using a milling tool, 1.5 mm cutting diameter (fig. 66 d), 18’000 rpm. (a) Fine milling of the cavital side in one level using a feed in the range from 0.05 to 0.13 mm, 1’100 mm/min feed rate. (b) Cavital side finished. (c) Fine milling of the occlusal side in one level, 0.1 mm feed, 1’100 mm/min feed rate. (d) Occlusal side finished. (Scale bars: 10 mm).
Following is an analysis of the milling process and its characteristics such as step sequence, types of milling, milling directions, cutting depth and alternatives. Concerning step sequence, roughing usually is carried out prior to fine machining in order to remove most of the material and to work out the rough shape of the framework. The sequence of roughing steps is not important due to allowance. In contrast it is observed that finishing of the cavital side prior to the occlusal side achieves better result at the cervical edges. A reason for this has not been found, so far.
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In most cases milling was performed with an alternating travel (see fig. 65) which in terms of milling technology means two different modes [43], conventional and climb milling. Edges created by both types of milling are shown in fig. 70. Better results are observed for conventional milling, especially the fabricated edge remains sharp in contrast to climb milling where chipping at the edges is observed. For conventional milling the tool feed and the cutting directions are opposed and therefore the cutting thickness changes from zero to its maximum. For climb milling the cutting thickness change is reversed for maximum to zero. A minimum chipping thickness at the cutting edge exit is required for a good cutting result without squeezing and friction effects [43]. Chipping is unacceptable especially at the thinnest parts of the framework with wall thickness in the range of 0.2 mm to 0.3 mm. Therefore we recommend to mill in conventional mode.
climb milling conventional milling Image: machining/types_milling_1.jpg
Fig. 70: Edge quality for conventional and climb milling type. A mill with ball end of 3 mm diameter was used with 10,000 rpm at 1,100 mm/min feed speed of to cut the groove in a cylindrical surface. Feed direction (of the mill) was from the right to the left side with increasing cutting depth from 0 to 6.0 mm. The conventional milled edge remains sharp (white arrows) and the climb milled edges show severe chippings (black arrows).
Cervical edges without chippings are achieved by rough milling with allowance of 0.4 mm and subsequent finish milling. The cervical edge of one of the cavities of a three-unit bridge scale-up by approximately 1.25 is shown in fig. 71. No break-outs are observed and the cervical edges are of high quality. Break-outs of small parts at the cervical edge are unacceptable and will lead to a rejection of the framework.
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Image: machining/edge_01.jpg
Fig. 71: Cervical edges after milling in presintered state. No break-out at the cervical edges is observed.
Chippings were already analysed prior in chapter IV.1.7. From the viewpoint of milling frameworks chippings of small size would be preferred. This assures that only small portions of the material are removed and that in case of unintentional break outs their size remains small.
Milling tool wear determines the endurance which is a function of the tool-to-materials interaction and which affects the accuracy and the surface quality. Endurance of PCD and cubic boron nitride (CBN) tools were found to be highly superior to HM, cermet and ceramic tools for turning green alumina compacts [46, 82]. These results may not be transferred to presintered TZP compacts due to the different microstructure of porous TZP blanks compared to porous alumina ceramics. Tool tips were inspected after milling TZP blanks. The tip of the 1.5 mm diameter HM milling tool after being in service for 60 crowns is shown in fig. 72. Typical wear characteristics on the major cutting edge, the flank or the clearance such as dull cutting edges, or crater wear are not visible. Wear of tools was only detected by a high-pitched sound during cutting. Expensive PCD
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and CBN tools benefit from their decrease in wear and hence for their increase in endurance. However, the cheaper HM tools exhibit a sufficient endurance of 100 crowns and therefore they were used in this work. A correction of the HM-tool diameter was not necessary due to its low wear characteristics.
(a)
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Image: machining/tool15_1.jpg
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Image: machining/tool15_2.jpg
Fig. 72: Inspection of the tool tip of a HM milling tool of 1.5 mm diameter after being in service for 60 crowns. (a) No wear is visible at the major cutting edge and the clearance and (b) at the chipping face.
Milling enlarged shapes has advantages for the machining process itself, the tooling, the part in comparison to machining in the original size. All details are scaled up especially the small ones. Hence, the machining process must not be as accurate as inaccuracy shrinks afterwards due to the sintering. Surfacial milling grooves which may act as strength-lowering flaws will also shrink during the final sintering. Tools with larger diameter may be used which possess more stability, less bending and less vibration, and therefore they enable higher feed rates. Enlarged parts also exhibit more stability and stiffness, and therefore withstand higher cutting forces. Machining may produce unintentionally created small cracks in the surfacial lay-
er. Their evidence during the green machining of alumina was reported by [46, 82]. Subsequent sintering may heals these cracks, however, neither their creation during milling of TZP blanks nor their healing were analysed in this work.
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Machining has been described for TZP blanks. A set of tools and parameters is provided to reliably mill frameworks out of blanks. The wear of HM tools is low and permits sufficient endurance. Hence, a small path has been established for an easy fabrication of complex surfaces by using well know machining techniques. Further work towards an optimization of the tool and the cutting parameters is required. Furthermore a machine principle has been developed (appendix: chapter VI.9.1), and experimental prototypes (appendix: fig. 126) built according to that principle proved to be feasible. The series machine for fabrication of ceramic frameworks is shown in fig. 127.
IV.3.5
Sintering: Achieving Net-Dimensions and Materials Properties
Introduction
Sintering is the final step for the fabrication of a TZP framework after it has been milled out of a blank. During sintering the porous framework consolidates and shrinks to the final dimensions. After sintering the part is ready to use and no additional machining in the sintered state is necessary. The sintering process is determined by its maximum temperature, the heating rate, the dwell time, the cooling rate and the atmosphere. The application of TZP in dentistry and in particular the DCM process require an isotropic shrinkage to maintain the quality and the accuracy of the cervical edge. Therefore the fit of the sintered frameworks on their affiliated master die was verified, and the shrinkage deviation from the prediction (see chapter IV.3.3) was determined by using a special calibration specimen. Good translucency of the framework material propagates the colour of the retainer teeth to the visible surface. This effect depends on the light transmission of the TZP material, and is highly desired by the dental technicians for aesthetic reasons.
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The aim of this chapter is to describe the sintering step for frameworks and to analyse the evolving properties of the TZP material. In particular we investigate the sintering process, microstructure, density, toughness, hardness, and translucency of sintered TZP. Furthermore the densification behaviour as function of sintering temperature and dwell time as well as its statistical variation are discussed. Accuracy during shrinkage is analysed using a calibration specimen and bridge on their respective master dies.
Materials and Methods
Dilatometry1 investigates the sintering behaviour of TZ-3YB and TZ-3YB-E to set-up the sinter process. We determined the start of sintering at 2 % shrinkage as a function of the compaction pressure and its affiliated density, maximum sintering rate, its temperature, and the end of sintering. Compacts of both materials were produced for various pressures in the range of 50 to 500 MPa, cut to a 2 mm x 2 mm x 8 mm cubic shape and then sintered in air atmosphere. Heating rate was 3 K/min to 1500 °C for TZ3YB and to 1400 °C for TZ-YB-E, respectively, with 120 min dwell time. Cooling was not controlled. Porous frameworks were sintered on flat, dense alumina plates in air under atmospheric pressure using two different high temperature resistance furnaces 2,3 each one equipped with Super Kanthal heating elements (see fig. 73). When sintering in the HT04-16E furnace2 a Silimantin muffle was used to diffuse the infrared radiation from the heating elements. No additional muffle in ETH-043 was necessary due to its special design. A small inertia mass of the refractories was chosen to enable high heating and cooling rates even though a higher refractory mass may produce a better temperature homogeneity inside the heating chamber. The sintering of restorations required a heating
1. 2. 3.
DIL 802S, Bähr Thermoanalyse GmbH, 32609 Hüllhorst, Germany Nabertherm HT 04-16E (Tony Güller AG, Hägendorf, Switzerland) ETH-04 (Nabertherm AG, Lilienthal / Bremen, Germany), prototype furnace
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rate of 1000 K/h up to 1500 °C for TZ-3YB and up to 1320 °C for TZ-3YB-E, respectively. Dwell time was 120 min for materials. The sintering temperature was chosen at values where small pellets1 sintered. Cooling was performed in the closed furnace to 150 °C before opening the furnace door. The cooling rate was not controlled.
(a)
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Image: Final_Sintering /furnace/ETH-04.jpg
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Fig. 73: High temperature resistance furnaces used for sintering. (a) prototype ETH04, (b) available HT04-16E.
For microstructural analysis pellets of 11 mm diameter and 3 mm height were sintered at 1500 °C for TZ-3YB and 1350 °C for TZ-3YB-E, respectively with 3 K/min heating rate and 120 min dwell time. Before polishing the pellets were glued to a steel plate and ground flat using a surface grinding machine2 equipped with a surface grinding wheel3. For polishing a lapping machine4 was used with a 1 µm diamond suspension5 for the final polishing step. Afterwards, the pellets were thermally etched for 30 min at temperatures 50 °C below the sintering temperature. After etching the pellets
1. 2. 3. 4. 5.
Cylindrical geometry of 11 mm diameter and 3 mm height in presintered state, density of 51 %TD. Chevalier FSG-818 AD, Falcon Machine Tools Co., Ltd., Taichung, Taiwan Diametal D54, concentration 75, binding B2, diameter 200 mm, width 9.0 mm, Diametal AG, Biel/Bienne, Switzerland FLM 300, A.W. Stähli AG, Läpptechnik, Pieterlen, Switzerland A.W. Stähli AG, Läpptechnik, Pieterlen, Switzerland
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were sputtered 1 with Au for 150 seconds under Ar atmosphere and analysed using SEM2. The grain size analysis was performed with the linear interception method [83]. We assumed spherical grains, and thus their equivalent grain diameter was calculated to 1.5 times of the mean intercept length. The Vickers hardness HV in MPa is determined according to chapter VI.2.3 (appendix) using eq. 16 (appendix), the calculation of the fracture toughness KIc in MPa√m is described in chapter VI.2.4 (appendix) using eq. 17, eq. 18 to eq. 19 (appendix). The experiments were carried out using polished, dense sintered samples prepared in the same way as for described microstructural analysis, but the thermal etching and the sputtering were omitted. The derivation of measurement error is found in chapter VI.3.4 and
chapter VI.3.5 (appendix) for the Vickers hardness and for the fracture toughness, respectively.
Translucency e.g. transmittance were investigated using dense sintered cylindrical plates made of TZ-3YB. The samples were sintered at 1500 °C with a heating rate of 3 K/min and 120 min dwell time, then they were ground to appropriate thickness and afterwards polished on both sides. They were 20 mm in diameter and had thicknesses in the range of 10 µm to 1’500 µm. Translucency was determined by two different methods. First the specimen was observed on a black background with a white stripe (contrast method) and second by the relative transmittance of light. The relative transmittance is measured by sending light from an Hg-lamp through the sample and then through an aperture of 3 mm in diameter and measuring the intensity behind the aperture with a lumimeter. As reference we measured the luminescence without a TZP plate in air between the lamp and the lumimeter. The measured values were compared with the
Bouguer-Lambert law (appendix: chapter VI.5) which theoretically describes the light transmittance through a material as a function of the sample thickness. The shrinkage homogeneity and the dimensional accuracy were analysed either by placing frameworks on their master cast or by the measurement of calibration specimen. Frameworks were fabricated, sintered, and cemented on their master gypsum cast. Then a mesial-distal cross-section was prepared and the gap between the cervical margin of the framework and the master die was analysed. This setup is shown in fig. 74. The cali1. 2.
Balzers SCD 40, Bal-Tec, 9496 Balzers, Principality of Liechtenstein Jeol JSM 6400
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bration specimen was fabricated by milling an array pattern with 3.0 mm x 3.0 mm quadratic elevations of 10.0 mm height out of a blank. The tolerances of the pattern in all spatial directions were better than 10 µm. The shape of the calibration specimen and the analysis method are shown in fig. 75. Before and after sintering the array pattern was measured recording distinct points. The data of its contour edges in presintered state was scaled down with the predicted shrinkage factor according to chapter IV.3.3, and then compared with the data of its contour edges in sintered state. Deviations of both data sets were computed. Straightness of the edges was analysed in the sintered state using a projector with a 20-fold magnification.
Image: final_sintering/bridge_accuracy_setup.jpg
Fig. 74: Dimensional accuracy and homogeneity of shrinkage determined using a three-unit bridge place on its master cast. Left: Front view. Right: Side view showing the mesial-distal cut.
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28 mm presintered specimen
Image: final_sintering/ calibration_specimen_02.jpg
Image: final_sintering/calibration_ ∆l) specimen_setup.wmfx ( Design: final_sintering/ l0 calibration_specimen1.wmf
prediction
calibration specimen
sintered specimen
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Fig. 75: Homogeneity of shrinkage and dimensional accuracy determined using a special calibration specimen. TZ-3YB, 1500 °C, 3 K/min heating rate, 120 min dwell time.
Process data were investigated by the density values of the fabricated frameworks. Statistical analysis includes 132 frameworks made of TZ-3YB and 30 made of TZ-3YBE, respectively. For density determination of the frameworks we used the Archimedes method. Three drops of liquid soap were added to the purified water for a better wetability of the samples. The relative error of the density is estimated at 0.1 % (appendix:
chapter VI.3.3).
Results and Discussion
The dilatometry of TZ-3YB (TZ-3YB-E) blanks showed the start of sintering at 1038 °C (1034 °C), the maximum sintering rate of -4.0 µ m/min (-7.3 µ m/min) at 1221 °C (1192 °C) , and the completion of sintering at 1500 °C (1400 °C) (fig. 76). The total linear shrinkage was found to be 22.9 % (21.3 %), the density of the specimen was 100 %TD. The shrinkage increased continuously with increasing temperature. The sintering rate showed only one maximum. Any shrinkage and expansion effects up to
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850 °C were already discussed (see chapter IV.1.6), therefore the range from RT to 850 °C was omitted in fig. 76. Hence, we conclude that full density is reliably achieved when sintering with a heating rate of 3 K/min up to 1500 °C with a dwell time of 120 min for TZ-3YB, and up to 1400 °C for TZ-3YB-E, respectively.
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Temperature [°C] Fig. 76: Shrinkage and sintering rate as a function of the temperature. (a) TZ-3YB, 1506 °C maximum temperature, (b) TZ-3YB-E, 1407 °C. Compaction pressure was 400 MPa, heating rate 3 K/min and dwell time 120 min at the maximum temperature.
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The sintering behaviour depends on the compaction pressure, e.g. the green density (fig. 77, fig. 78) for TZ-3YB and TZ-3YB-E. The density of all sintered samples exceeded 99 %TD. Increasing compaction pressure shifted the start of the sintering to higher temperatures and the total linear shrinkage to lower values. The maximum sintering rate remained constant for TZ-3YB, and a slight decrease was found for TZ-3YB-E. The extrapolation of the sintering ability to larger sized badies such as big dental restorations is limited, as we used small specimen in this investigation. Initial compaction pressure has a major effect on the sintering process. For all investigated compaction pressures from 50 MPa to 500 MPa, e.g. blank densities from 2.5 g/cm3 to 3.2 g/cm3 sintering lead to full density. Hence, we conclude that a 3 K/min heating rate up to 1500 °C for TZ-3YB and up to 1400 °C for TZ-3YB-E and a 120 min dwell time at maximum temperature is appropriate for further investigations.
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Compaction Pressure [MPa] Fig. 77: Characteristics of sintering process for TZ-3YB and TZ-3YB-E blanks as a function of the compaction pressure. (a) Start of sintering at -2 % shrinkage. (b) Maximum sinter rate.
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Compaction Pressure [MPa] Fig. 78: Characteristics of sintering process for TZ-3YB and TZ-3YB-E blanks as a function of the compaction pressure. (b) Temperature at the maximum sintering rate. (b) Total linear shrinkage.
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The microstructures of sintered TZP showed a dense, fine polycrystaline structure with a homogeneous grain size distribution. The SEM pictures are shown in fig. 79 and
fig. 80 for TZ-3YB and for TZ-3YB-E, respectively. The grain size distribution was narrow: the large ones measured approximately 0.4 µm, the small ones about 0.1 µm. The shape of the grains was equiaxed, and some of them exhibited a pentagonal shape. Exaggerated growth of single grains, isotropy, texture or secondary phases were not detected. The microstructure of TZ-3YB displays smaller grain sizes and more homogeneous shapes compared to TZ-3YB-E. For both TZP’s we found a submircon-sized microstructure in which no flaws were observed. We therefore expect good mechanical properties.
Image:: Final_Sintering/microstructure _TZP/TZP-micro.jpg Design: Final_Sintering/microstructure _TZP/TZP-micro.psd
Fig. 79: Microstructure of sintered TZ-3YB, 1500 °C, 3 K/min heating rate, 120 min dwell time.
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Image:: Final_Sintering/microstructure_E-TZP/E-TZP-micro.jpg Design: Final_Sintering/microstructure_E-TZPE/E-TZP-micro.psd
Fig. 80: Microstructure of sintered TZ-3YB-E, 1350 °C, 3 K/min heating rate, 120 min dwell time.
The quantitative analysis of the microstructures confirms the homogeneous grain size distribution for both TZP materials as seen in their micrographs. The results of the grain statistics by means of the line-intercept-method are shown in fig. 81. In the case of TZ-3YB the grain size ranges from less than 0.1 µm to 1 µm with a median grain diameter d50 of 0.32 µm. TZ-3YB-E grains range from 0.05 µm to 0.7 µm in diameter with a median grain diameter d50 of 0.15 µm which is less than half the size compared to that of TZ-3YB. In comparison the grains of the TZ-3YB-E are smaller sized than those of the TZ-3YB material.
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Image: microstructure_TZP/TZP_microstructure-plot.WMF Kaleidagraph Data: microstructure_TZP/TZP_microstructure-plot.qpc 40 Design: microstructure_TZP/microstructure-TZP.ppt
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Grain Diameter [nm] Fig. 81: Grain size analysis for (a) TZ-3YB sintered at 1500 °C, and (b) TZ-YB-E sintered at 1350 °C. Heating rate 3 K/min, dwell time 120 min.
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The microstructures of the 3 mol-% yttria stabilized tetragonal zirconia (TZ-3Y)1-
alumina-silica system has already been investigated for varying alumina content from 0 to 0.89 wt-%, and silica content from 0 to 0.8 wt-%, respectively [18]. Secondary needle-like alumina-rich phases were found in the TZ-3Y with 0.89 wt-% alumina as well as grain-sized pores in the pure TZ-3Y and the TZ-3Y with 0.28 mol-% alumina, respectively. Secondary phases or pores are not desired. The grain size of TZ-3Y is similar to that of TZ-3YB used in this work, however the TZ-3Y with addition of 0.28 wt-% alumina possesses a grain size that is three times coarser than that of TZ-3YB-E used this work. However, a fine grained pore-free microstructure is desired for high strength and high reliability as required in dental application. That TZ-3Y with alumina additions fabricated by Michel [18], although it shows a similar composition than TZ-3YB-E, is not suited for an application in dentistry. The Vickers hardness in MPa as a function of the applied load in N is shown in fig.
82. Loads up to 30 N lead to dropping hardness values, whereas for higher loads the hardness stayed constant. The hardness we found is slightly higher than the 1250 MPa given by the powder manufacturer [34] . Values in the 10 to 13 GPa range were reported using different processing routes [18, 84, 85], and however they correspond well to our findings if divided by 10 2. The toughness or the critical stress intensity factor KIc results calculated according to the formula given by Niihara et. al. and given by Anstis et. al. (both appendix: chapter
VI.2.4) are shown fig. 82. They were in the range between 5 MPa√m to 8 MPa√m for the applied loads and for the calculation methods. Typical Palmqvist radial cracks due to Vickers indentation for appropriate high loads as taken for the KIc analysis are shown in
fig. 83. The KIc values decreased with increasing loads, and approached 4 MPa√m to 5 MPa√m. Both evaluation methods show the same behaviour, however Anstis-values were about l MPa√m below those calculated with the Niihara-formula. Extraordinarly high toughness of 15 MPa√m for TZP were reported by Tsukuma et. al. [87] and attributed to the beneficial transformation toughening mechanism found for alloyed zirconias (see for example [88]). Ruiz and Ready [84] found the toughness to increase with higher 1. 2.
The 3 mol-% yttria stabilized tetragonal zirconia which was fabricated by Michel [18] is termed TZ-3Y to distinguish it from Tosoh’s TZP materials. Values 10 times higher indicate that the authors used the prior HV practice and ignored to divide by 9.81 when using load in Newton (according to [86]).
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sintering temperature starting from about 4 MPa √m for 1450 °C to 10 MPa √m for 1750 °C, respectively. Alloying TZP with alumina was reported to reduce its toughness drastically to about 6 MPa√m [87]. This is not in agreement with the technical data of TZ-3YB-E as supplied by the manufacturer reporting equal toughness of 5 MPa√m for both their TZPs [34, 35]. The values determined here are in agreement with literature data additionally considering that the maximum sintering temperature was 1500 °C. In order to achieve a high toughness a high sintering temperature is be recommended, how-
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4 2
Toughness KIc [MPa √m]
Vickers Hardness [MPa]
ever this is usually in contrast to achieving a small grain size.
0
Load [N] Fig. 82: Vickers hardness and fracture toughness of TZ-3YB by indentation as function of the applied load. For error calculation see chapter VI.3.4 (appendix) and chapter VI.3.5 (appendix) for the hardness and for the fracture toughness, respectively. Sintering 1500 °C, heating rate 3 K/min, dwell time 120 min, samples polished.
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Fig. 83: Vickers indentation in TZP showing the Palmqvist cracks (blank arrows). Sintering 1500 °C, heating rate 3 K/min, dwell time 120 min, samples polished.
In the following we investigate long dwell times, various sintering temperatures and different heating rates on the TZP’s materials properties. These process parameters are discussed as they affect the evolving density, phase content, grain size, hardness and the fracture toughness. The density of TZ-3YB as a function dwell times up to 700 hours and different sintering temperatures of 1450°C, 1500 °C and 1550°C is shown in fig. 84. All samples achieved 100 %TD (6.05 g/cm3), however to achieve their full density it took different dwell times depending on the temperature. Higher maximum temperatures promoted a shorter dwell time and vice versa. Long dwell times cause the density to decrease, it declines earlier and faster with higher temperatures. For holding times longer than 200 h the sample’s density drops below 99 %TD for 1500 °C and 1550°C.
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6.1
6.0
99
Image: figures/dwell_time_temperature_density.wmf Design: figures/dwell_time_temperature_density.ppt Origin Data: figures/dens_tem.opj 5.9
98
1450 °C 1500 °C 1550 °C 5.8
1
10
100
97
Rel. Density [%TD]
Density [g/cm3]
100
1000
Dwell Time [h] Fig. 84: Influence of the maximum sintering temperature and the dwell time on the evolving density for TZ-3YB [17]. The relative error of the density is estimated to be smaller than 1.0 %. Heating rate was 10 K/h up to 500 °C and 100 K/ h up to the sintering temperature, respectively. Samples were quenched in air.
Changes in phase content for TZ-3YB were not observed at 1500 °C for dwell times up to 150 h [17]. However, for longer dwell times the content of tetragonal phase was reduced, whereas the monoclinic and the cubic fraction increased. Higher sintering temperatures lead to a decrease of the tetragonal phase and simultaneously to an increase of cubic and monoclinic content [84]. As the tetragonal phase is desired due to its inherent toughening mechanism we recommend sintering temperatures lower than 1500 °C and dwell times shorter than 150 h. The grain size generally becomes larger as the dwell time or the sintering temperature is increased [17, 84, 89]. Its susceptibility to the prolonged holding time was lower than to the sintering temperature [17]. Grain growth kinetics were found to be much slower for the tetragonal phase1 (TZP) compared to other ceramic materials (alumina) and compared to its cubic and its monoclinic zirconia phases [90, 91, 94]. It also be-
1.
Grain growth factors of TZP: 0.02 x 10-22 m3/s at 1400 °C to 0.49 x 10-22 m3/s at 1550 °C [89]
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comes much faster with rising sintering temperatures1 [89, 91]. From the preceding argumentation we recommend a sintering temperature of 1500 °C for TZ-3YB at a minimal possible dwell time to allow firing to full density and simultaneously achieving a small grain size. The hardness of TZP was found to decrease slightly from initially about 12 GPa to 11 GPa for dwell times up to 400 h at a constant sintering temperature of 1500 °C [17]. In the range between 1450 °C to 1750 °C and a constant 2 h dwell time a similar decline of the hardness also was reported independently by Ruiz and Readey [84]. The investigation of the fracture toughness of TZ-3YB proved a high sensitivity to the dwell time and the temperature. Using the Evans-formula [92] the fracture toughness values started at 5.6 MPa√m for 1 h dwell time at 1500 °C, reached its maximum values of 11 to 13 MPa√m for more than 100 h dwell time, and then decreased to 3.5 MPa√m for more than 520 h dwell time, respectively [17]. For temperatures in the range between 1450 °C and 1750 °C and 2 h dwell time the fracture toughness was determined to rise from 5 MPa√m to 10 MPa√m, respectively [84].
Heating rates in the range between 2 K/min and 200 K/min and 120 min dwell time had a negligible influence on the density and on the grain size of TZ-3YB [38, 93, 94]. For comparison we achieved a maximum heating rate of 17 K/min with our resistance furnaces. However in the case of TZP high heating rates involve the risk of consolidating the outer surface before the inner bulk, and thus hindering the inner densification or the risk of entrapped gas [93] both resulting in a poor microstructure. From these discussions our conclusion is that the TZ-3YB material behaves tolerantly when applying extreme sintering conditions. However, the sintering parameters have to be adjusted in compromise with each other to best produce the required materials properties. For TZ-3YB we recommend sintering at 1500 °C with a short dwell time of less than 8 hours, at best 120 min. All heating rates are possible in the range from 1 K/min to 17 K/min. With this set of parameters a reliable fabrication of dental restorations should be possible. We assume that similar results are achieved for the TZ-3YB-E and therefore that conclusions apply equal to those for the TZ-3YB material. However the sintering temperature should be 1350 °C.
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The translucency or transmittance of sintered TZP samples for various thicknesses is shown in fig. 85 (contrast test). With increasing thickness of the plates the translucency decreases as expected. The white stripe could be recognized for the 0.78 mm sample thickness which is more than the frameworks’ usual thickness at the cavities. The transmittance of white light as a function of the sample’s thickness is shown in fig. 86. It decreases rapidly with increasing thickness. For 1 mm thickness only about 4 % of the light is transmitted, for 0.3 mm approximately 7 % is measured. The visual appearance of the contrast test and the measurement of transmittance correspond to each other. Transmittance of monochromatic light can be theoretically described by the Bouguer-
Lambert law (appendix: chapter VI.5). Its exponential decay is in good qualitative correspondence to our measurements and fits well for the thin samples despite that their deviation grows for thicknesses larger than 200 µm. However a higher number of measurements should be performed for a better curve fitting. Compared to opaque metal frameworks which are standard of excellence for bridges, TZP exhibits translucency which enables the dental technician to achieve better aesthetic results. However, TZP materials show inferior light transmittance when compared to highly translucent glass-ceramics such as Empress.
Image: Final_Sintering\translucency.jpg Design: Final_Sintering\translucency.psd
Fig. 85: Translucency of TZ-3YB as a function of samples thickness.
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Transmittance [%]
60
178
Transmittance Image: Final_Sintering\transmittance.wmf Bouguer-Lambert (Fit) Design: Final_Sintering\transmittance.ppt A: 66.985; B: 100.353 Origin Data: Final_Sintering\transmittance.opj
50 40 30 20 10 0 0
200
400
600
800 1000 1200 1400 1600
Thickness [µm] Fig. 86: Transmittance as a function of the TZ-3YB sample thickness. The fit is based on the Bouguer-Lambert law (appendix: eq. 32) [95].
Dimensional accuracy and shrinkage homogeneity were determined for the calibration specimen and for frameworks, respectively. The calibration specimen as shown in fig. 75 possessed a homogeneous linear sintering shrinkage of 22.807 % in all three spatial directions. During sintering the edges of the elevated posts remained straight and did not bulge. The edges remained sharp and did not show significant rounding in contrast to our expectations. We assumed rounding effects would occur due to material transport during sintering [7]. The average deviation between the predicted dimensions and the sintered array pattern was an absolute 19 µm ± 12 µm standard deviation for 96 measuring points on the calibration specimen. The average was detected to be equal in all spatial directions and showed no significant deviation. The relative dimensional accuracy is therefore better than 0.07 % on the 28 mm total length of the array pattern in the presintered state or better than 0.09 % if referenced to the length in sintered state. Hence the accuracy after sintering is proved to be sufficient for the dental application, and no subsequent hard machining will be needed afterwards.
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The cross-section of a three-unit bridge on its master cast is shown in fig. 87. A 80 µm gap at the cervical edge is found which slightly exceeds the required tolerances of 50 µ m. The inner cavity of the framework does not follow the master cast exactly in 50 µm parallel distance. To its majority this was found to be due to the wax model of the framework. However, this framework analysis reveals the accuracy over the complete DCM process. Many influences such as manual work and post-cementation co-determine the gap distance. System inherent errors constrain the probability for matching the required 50 µm tolerance, even if the fabrication is performed precisely by skilled dental technicians [74]. A long span five-unit complex shaped framework with extreme cranio-caudal curvature and three abutments is shown in fig. 88. This type of work is placed towards the upper level of complexity due to its long span, the multiple abutments and its curvature. Nevertheless its fitting on its master cast proves a sufficient dimensional accuracy and shrinkage homogeneity. Therefore we conclude that the DCM process is feasible for its application in dentistry from the viewpoint of dimensional accuracy and tolerance.
veneer porcelain
Image: Final_Sintering\bridge_accuracy_02.jpg
master cast framework cement
Fig. 87: Dimensional accuracy and shrinkage homogeneity of sintered three-unit framework. Zoom-in from left side to right side. The framework fits on the master cast and shows a cementation gap at the cervical margin of 80 µm.
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(a)
180
(b)
Image: Final_Sintering\5-bridge_2.jpg
Image: Final_Sintering\5-bridge_1.jpg
Fig. 88: Example of a five-unit framework. (a) Enlarged framework and sintered framework demonstrate the amount of shrinkage. (b) Sintered framework placed on its master gypsum model proves its fit.
The variation of the final density was investigated by the means of statistical analysis for fabricated frameworks. The sintered densities of TZ-3YB and TZ-3YB-E are shown as a box plot in fig. 89. The mean density was determined at 6.07 g/cm3 for TZ3YB and at 6.05 g/cm3 for TZ-3YB-E, respectively. In the case of TZ-3YB-E our initial assumption was accurate for the final density, and no further refinement has therefore been necessary to calculate the enlargement factor. For TZ-3YB the mean final density was 0.3 % higher than our initial assumption, and therefore has to be refined. Variation from minimum to maximum density value was ±1 % for TZ-3YB and ±0.6 % for TZ3YB-E, respectively. Error in measuring accounts for only 0.1 % and hence is not responsible for that variation. However, milling powder remnant which was baked on the frameworks’ surface during sintering is claimed responsible for most of that variation. This remnant is not dense and therefore stores air. It shows a non-negligible influence on
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the frameworks’ density and the density measurement error, respectively. Variation exceeds the required maximum tolerance of 0.17 %1 (see eq. 13). Despite this, the fabricated and sintered frameworks using our initial assumptions did not show any misfit. Therefore, we conclude that density must not necessarily be refined for both materials.
Density [g/cm3]
6.15
6.10
6.05
Image: final_sintering/process/process.wmf Design: final_sintering/process/process.ppt Data: final_sintering/process/process.opj
6.00
5.95
n= 132
n= 30
TZ-3YB
TZ-3YB-E
Fig. 89: Variation of density of sintered frameworks. Minimal value (triangle), mean value (square), maximum value (circle). The box presents the 5 th, 25 th, 50 th (median), 75 th and 95 th percentile.
δls 1 p + δρ ------- = δ----f = --- ⋅ δρ ----------------s- ≤ 0.17% ls 3 ρp ρs f
1.
Eq. 13
Tolerance is 0.17 % under the assumptions first that δls is smaller than 50 µm on a sintered length ls of 30 mm, second that the presintered density’s error is negligible, and third that the milling tolerance is zero (see eq. 13)
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Summary
Sintering is the last step in the fabrication of the framework. During this step the materials properties emerge and the framework shrinks to its original dimensions. In the preceding chapter we determined the process parameters for sintering: the temperature, the heating rate and the dwell time for TZ-3YB and for TZ-3YB-E, respectively. Both TZP materials exhibit a similar densification behaviour during sintering. However the major differences were found to be the sintering temperature and the shrinkage rate. With increasing green density we found that the starting temperature for sintering is shifted to higher values, the shrinkage decreases, as well as the temperature of maximum shrinkage rate. In the case of TZ-3YB the sintering rate stays constant for all green densities whereas in the case of TZ-3YB-E it is slightly increased. The microstructure for both TZPs was homogeneous, dense, and of submicron grain size. However, the TZ-3YB-E possesses a 150 µm grain size which is more than two times smaller than that of TZ-3YB (322 µ m). After sintering we determined the
Vickers hardness at more than 1’300 MPa and the toughness to be higher than 4 MPa√m. These values are in good agreement with literature datae when the sintering parameters are considered, it is however possible to achieve higher toughness. The dwell time and the sintering temperature affects the density, the grain size, hardness and toughness. Full density was achieved at temperatures in the range of 1450 °C to 1550 °C for a dwell time of up to 8 h. Dwell times over 100 h resulted in a decrease of density. Phase changes were not observed for sintering times up to 150 h, although prolonged sintering is responsible for the reduction of the tetragonal phase content. The grains only grow very slowly and therefore the fine microstructure is preserved over a large range of sintering times. Vickers Hardness in the range from 11 GPa to 12 GPa exhibits minor sensitivity to a variation of both sintering parameters (see page 172), whereas toughness is particularly susceptible having only 5.6 MP√m for 1500 °C with 1 h dwell time. It increased significantly up to 10 MP√m or 13 MP√m when sintered at 1750 °C for 1 hour or 1500 °C for 100 h, respectively.
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The heating rates we could apply at our facility did not change the final density or the densification behaviour nor did they affect the microstructure of TZP. Much higher heating rates resulted in the same findings, but for bigger parts involve the risk of sintering the outer surface of a part which hinders the inner shrinkage. This may lead to a poor microstructure of the bulk material. From the preceding results we concluded the sintering parameters and recommend to consolidate the TZ-3YB at 1500 °C and the TZ-3YB-E at 1350 °C, respectively, and to use a heating rate of 1000 K/h with 120 min dwell time. TZP material as processed here possesses a certain translucency as shown by contrast testing. With increasing sample thickness the transmittance declines, the measured values match qualitatively the theoretic Bouguer-Lambert law. For thickness of 0.2 mm we determined the transmittance of TZ-3YB to be 10 % and for 0.8 mm to be 5 %. This light transmittance of the framework material enables the dental technician to benefit from the natural colour of the retainers and hence to produce better aesthetic results. The absolute dimensional accuracy was determined by means of an array shaped calibration specimen to 19 µm ± 12 µm and the relative tolerance was found to be better than 0.07 %. This is sufficient for the application of the DCM in dentistry. Furthermore the shrinkage in all three spatial directions were proven to be isotropic. Even complexly shaped long-spanning frameworks showed a good fit on their master cast models after final sintering. Therefore we conclude that the DCM process is feasible in dental application. The cross-section through a bridge showed an acceptable marginal fit and the inner cavity follows well the model contours. Adequate sintering parameters were established for dental restorations made of TZ-3YB and TZ-3YB-E. Sintering of frameworks and simultaneously achieving the dimensional accuracy after sintering has been demonstrated. Initial parameter sets were established, but further investigation need to be carried out.
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IV.3.6
184
Veneering: Coating the Framework
Introduction
Coating the zirconia framework with a veneer porcelain is the last fabrication step for a restoration. The reasons for applying a veneer porcelain are two-fold. First, each patient has an individual non-uniform tooth colour. Second, the antagonistic teeth may be destroyed by TZP material due to its hardness and abrasivity. The veneer porcelain is coated on the TZP framework in an artistic manner in multiple layers of different colours. Outward shape is formed according to the natural tooth. The aim of the veneering step is to adjust appearance of artificial restorations as close as possible to the natural tooth. This chapter will describe the procedure of applying the veneer porcelain on the densely sintered TZP framework. The veneering work was done at the Zentrum für
Zahn-, Mund-, und Kiefernheilkunde der Universität Zürich, Klinik für Kronen-BrückeProthetik, Teilprothetik und Zahnärztliche Materialkunde by experienced dental technicians. The experimental veneer porcelain W35/11 described in chapter IV.2 was used. Material and Methods
W35/11 powder1 (see chapter IV.2) is mixed with water or with a special model aqueous agent2 to a pasty consistency on a glass plate. Then the paste is applied to the framework using red marten brush. The firing is performed in a Ivoclar P95 furnace2 according to tab. 12.
1. 2.
A set of colours using W35/11 base material was supplied by Ivoclar, Schaan Principality of Liechtenstein. Ivoclar, Schaan Principality of Liechtenstein
185
cycle
stand-by heating temp. rate [°C] [K/min]
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closing time [min.]
dwell time [min.]
firing temp. [°C]
vacuum application temp. [°C]
first dentin firing
400
60
4
1
890
580
second dentin firing
400
60
4
1
880
580
glazing (pure)
400
60
4
0.4
910
no
glazing (colour)
400
60
4
2
870
no
Tab. 12: Firing cycles for applying the veneer porcelain on TZP frameworks.
The coating of the zirconia framework requires to application of multiple porcelain layers of different colours. This procedure is shown in fig. 90. Each layer is painted using a brush. Then it is dried, the excess material is removed using a spatula and afterwards fired on a mesh-tray in the furnace at temperatures between 870 °C to 910 °C for 0.4 min to 2 min (for details see tab. 12). Before applying a next layer, the last one is ground in order to refine the shape of the restoration and to roughen the surface. This process is performed serval times until the dental technician is satisfied with its shape and its aesthetic appearance. Then, the glazing is performed as the last step in order to close all the open surfice porosity and to give the surface its bright visual shine.
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(a)
(b)
(c)
(d)
(e)
(f)
(g)
(h) Image: coating/veneer-1.jpg
Fig. 90: Coating the zirconia framework with veneer porcelain (Courtesy of the University of Zurich). (a) Initial framework. (b) Painting the pasty veneer porcelain slurry. (c) Removing excess material with a spatula. (d) Firing the porcelain onto the framework. (e) Intermediate state after firing from the lingual view. (f) The same from the vestibulary view. (g) Grinding the porcelain. (h) Finished framework with veneer porcelain, ready for use.
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Results and Discussion
In this chapter the feasibility of W35/11 porcelain has been proven for complex framework shapes which is a precondition for its application in clinical studies. Out of the scope of this work are extended pre-clinical tests usually performed by dental material suppliers such as shear strength, fatigue, and thermocycling between 5 °C - 37 °C 55 °C. Within further clinical tests the veneer porcelain has to prove its reliability. The aesthetic result and the ease of handling of W35/11 veneer porcelain proved to be acceptable for dental technicians using the firing cycles listed in tab. 12. But, the available colour palette provided by Ivoclar for this work limit the aesthetical possibilities. Stability of the veneer layers enabled the use of the layering concept during firing, and is one precondition for highest aesthetical results later. However, further improvement of the handling properties may be necessary.
IV.3.7
Summary
A new process called Direct Ceramic Machining (DCM) for the fabrication of allceramic dental restorations has been developed. It mills a TZP framework out of a porous, easily machinable TZP blank and subsequently sinters it to full density. Then it is coated with a porcelain veneer using conventional dental methods to finish the restoration. The framework’s shape is digitally acquired from a plastic model by means of mechanical tactile or of optical technologies. Afterwards the shape data is enlarged to compensate for the sintering shrinkage, and then tool path information is derived from the enlarged shape to allow controlling a three axis milling machine. The preceding subchapters investigated the DCM process steps and established a small pathway to prove the feasibility. Different tools and parameters were evaluated. For this we used blanks either made from TZ-3YB powder or from TZ-3YB-E, respectively (see chapter IV.1).
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Two different tactile and one optical conoscopical system were successfully applied for digitizing. Tactile technology had advantages at sharp edges, undercuts, steep surfaces and its shape-adapted data density leading to small file sizes. It showed a good accuracy depending on the scanning parameters, however the contact-to-surface causes deformation of the models and debris limits its ability to scan wax models. The optical technology had its benefits in the non-contact working, inertia-freeness and the small beam diameter which enables to measure without deforming the model surface small inner cavities at high speeds. In case of tactile digitizing we achieved best results for plastic frameworks with a stylus of 1.2 mm diameter at speeds lower than 1000 mm/min. For the cavital side we used variable feed between 0.05 mm and 0.13 mm, and for the occlusal 0.13 mm, respectively. The optical system allowed to scan wax models with up to 2.4 m/min speed and a frequency of data acquisition of 800 Hz. The result of the scans are unconnected points of the frameworks surface, so called point clouds. In case of tactile the mean point density was about 160 dpi and of optical it was 800 dpi, both were sufficient to produce good accuracy. The enlargement uses the point clouds and homogeneously scales them up by approximately 25 % in all spatial directions. The model is based on an afine transformation with one enlargement factor which is determined by its macroscopic and therefore easyto-measure relative density in %TD. However, this method of enlargement relies on blanks shrinking isotropically in order to match the final dimensions. It also depends on the a priori final density after firing for which we had to make an reasonable assumption. We assumed to sinter to 100 %TD. Blanks’ fabrication tolerances produced a variation of its density which required to determine the enlargement factor for each individually. To avoid this expense its tolerance has to be smaller than 0.17 % and the relative errors of the remaining DCM process steps must be negligible. The machining cuts an enlarged framework out of a porous ceramic blank with the tool path information generated using the scaled-up point clouds. During this work we established an experimental setup for machining combining the machine, tools, clamping device, sequence for milling and a parameter set well suited for frameworks. Various milling tools of different shapes and materials were investigated, and at the end we recommend HM mills with 1.2 to 1.5 mm diameter and 17 mm cutting length for finishing and 2.5 to 3.0 mm diameter and 20 mm cutting length for roughing. HM tools showed a
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sufficient life-time for more than 60 crowns without visible wear. Best milling results with high quality gracile cervical edges and absence of unintentional break-outs were achieved for a cutting speed of 1.5 m/s and feed rates of up to 2 m/min, however conventional milling mode was required. As this step uses high precision machining devices and tools we assume that the error of this step is negligible in comparison to the other DCM process steps. During the sintering which is the last step in the fabrication of a framework, the materials properties evolve and the framework shrinks to its net-dimensions. We investigated the sintering temperature, the heating rate and the dwell time and established an adequate parameter set for both TZP materials: 1000 K/h heating rate, 120 min dwell time at 1500 °C for TZ-3YB and at 1350 °C for TZ-3YB-E, respectively. The densification behaviour for both materials was similar except for the sintering temperature, total shrinkage and the shrinkage rate. For various green densities changes of the starting temperature, total shrinkage and the temperature of maximum sintering rate were observed whereas the sintering temperature, rate and the end stayed constant. A fine submicronsized homogeneous and dense mircrostructure with a median grain size of 322 nm were determined for TZ-3YB and of 150 nm for TZ-3YB-E, respectively. Hardness between 11 to 13 GPa and toughness from 5 to 12 MPa√m were determined in accordance with literature data. Long dwell times showed a strong influence on density and toughness. After sintering the frameworks showed a beneficial translucency. Dimensional accuracy was quantified to be 19 µm ± 12 µm absolute and 0.07 % relative. Shrinkage was isotropic in all spatial directions. In combination with the good fit of all frameworks when placed on their master dies we conclude that the dimensional accuracy is sufficient for dentistry. A final density variation was found but we attributed this to the powder remnants which originated during milling and which were not cleaned before sintering. Therefore dimensional accuracy was not impaired. The veneering of the TZP frameworks is performed using conventional dental methods. For veneer material we used the porcelain described in chapter IV.2.
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During the work we fabricated 132 restorations from TZ-3YB and 30 from TZ3YB-E, respectively, all had a sufficient accuracy. The core requirements like accuracy, reliable fabrication, simple handling (in principle) were demonstrated. Hence the DCM
process proved to be feasible for application in dentistry. Nevertheless, further investigations for improvements are necessary and the limits of the process remain to be researched.
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IV.4
Preclinical Performance of the Bridges: Load Bearing Capacity and Reliability
IV.4.1
Introduction
TZP shows a much higher bend strength and stress intensity factor KIc compared to alumina or other porcelains or glass-ceramics (see fig. 1). Despite its favourable mechanical materials properties, TZP has not yet been proven to bear loads for posterior all-ceramic bridges in the oral environment. Hence, preclinical analysis which takes into account simulataneously the mechanical loads in the oral environement and the clinical failure modes is required prior to performing any clinicial investigation.
Clinical fractures of all-ceramic crowns and bridges have rarely been identified. Crowns were reported to fail from the cavital cementation surface which is opposed to the chewing surface [96-98] whereas all-ceramic bridges failed at their connectors [99]. A majority of the cracks originated at the interface of framework to veneer porcelain. A typical fracture of a anterior three-unit all-ceramic bridge in the upper jaw is shown in
fig. 91. Kelly [96] claimed that some in vitro investigations suffered from clinical nonrelevant fractures such as Hertzian contact stress fractures. The analysis with immobile retainers produced a two to three times higher failure load than found when testing with mobile posts [100], and therefore lead to the overestimation of the load bearing capacity of the material and its clinical possibilities, respectively. Hence, a preclinical load bearing analysis must be guided by those experiences and produce clinically relevant fractures in the interdental connector with initiation from the gingival side.
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image: load_bearing_of_bridges /clinical_gesamt.jpg
Fig. 91: Clinical fracture observed for an all-ceramic anterior bridge in the upper jaw (Courtesy: University of Zurich). White arrows show the fracture from the facial view (top left) and from the lingual view (bottom left). Top right shows the retainer from the lateral view, and bottom right the pontic from lateral view, respectively.
This chapter investigates the static load bearing capacity and the reliability of TZP frameworks fabricated by the DCM process in comparison to other commercially available all-ceramic dental systems such as Vita Celay In-Ceram Alumina and IPS Empress2. Furthermore it also aims to analyse the influence of the veneer porcelain on the load bearing capacity and the reliabilty. With the probable mastication forces in mind, a FEA and a load bearing capacity analysis provide the means for the estimation of the bridges failure probability and therefore for giving recommendations on the connector dimensions.
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IV.4.2
Direct Ceramic Machining Process
Material and Methods
Three-unit frameworks of TZ-3YB, TZ-3YB-E, In-Ceram Alumina, and IPS Empress2 were fabricated. The TZ-3YB and TZ-3YB-E frameworks were fabricated by the DCM process as described in chapter IV.3. In-Ceram frameworks were fabricated according to the manufacturer’s instructions by an experienced dental technician. Prefabricated Vita Alumina 1 blanks were ground on a manually operated Celay 2 copy grinding machine and afterwards infiltrated with an appropriate lanthanum glass. IPS Empress2 frameworks were fabricated according also to the manufacturer’s instructions3. The TZP was veneered using the experimental W35/11 porcelain, and the In-Ceram using Vitadur Alpha4. The shape of the frameworks were reproduced from two master models with different distal connector sizes (see fig. 92): one had a 6.9 mm2 average cross-section area 5 (occlusogingival height 2.7 mm, buccolingual width 2.6 mm), and the other one of 10.6 mm2 (occlusogingival height 3.2 mm, buccolingual width 3.3 mm), respectively. The mesial connector cross-section area was always 1 mm2 larger than the distal one, exhibiting 7.9 mm2 and 11.6 mm2, respectively. The design of the frameworks simulated a three-unit bridge from the second premolar (tooth no. 5) to the second molar (tooth no. 7). It was simplified by flattening its occlusal surface. All frameworks were equally shaped at the abutement teeth: 6° convergence at the inner cavity, 5 mm total height, 0.8 mm occlusal wall thickness and 1.0 mm lateral wall thickness. It’s pontics also exhibited similar shape. In case of TZ-3YB, IPS Empress2 and In-Ceram the flat occlusal surface was produced by grinding and afterwards polishing 6 , in case of TZ-3YB-E grinding was omitted.
1. 2. 3. 4. 5. 6.
Vita Zahnfabrik, Bad Säckingen, Germany Mikrona, Spreitenbach, Switzerland Ivoclar, Schaan, Principality of Liechtenstein Vita Zahnfabrik, Bad Säckingen, Germany Cross-section area was measured using a calliper rule at the thinnest part of connector and calculating using formula for rectangles. Accutom, Struers GmbH, Birmensdorf, Switzerland
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194
image: load_bearing_of_bridges /framework/framework_2.jpg
height
width
image: load_bearing_of_bridges /cross-section.jpg
Fig. 92: Design of frameworks for load bearing and reliability investitagions. Left side: three-unit framework made of TZP. Right side: View of the cross-section of a 10.5 mm2 interdental connector (3.4 mm buccolingual width, 3.1 mm occlusogingival height).
Different veneer porcelain thicknesses of 0.5 mm, 1.0 mm and 1.5 mm at the gingival side of the connectors were analysed for the 6.9 mm2 frameworks. The thickness was checked by light microscopy after its fracture (see fig. 93). The W35/11 veneer was applied in thicknesses of 0.5 mm, 1.0 mm and 1.5 mm with three, four, three firings, respectively, the first one at 930 °C and the subsequent ones at 920 °C. The In-Ceram samples were veneered with Vitadur Alpha1 in three firings according to the manufacturer’s instructions. An overview of the test sets including the fabricated amount of specimen is presented in tab. 13.
1.
Vita Zahnfabrik, Bad Säckingen, Germany
195
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Image: load_bearing_of_bridges/porcelain/porcelain_gesamt.jpg
Fig. 93: W35/11 porcelain thickness (grey) on TZP framework (white) (Courtesy: of University of Zurich). Left: 0.56 mm measured thickness, 0.5 mm target. Middle: 0.8 mm measured, 1.0 mm target. Right: 1.4 mm to 1.5 mm measured, 1.5 mm target.
material
6.9 mm2 connector
10.6 mm2 connector
0.5 mm veneer
1.0 mm veneer
1.5 mm veneer
TZ-3YB
15
15
3
3
15
TZ-3Y-E
15
-
-
-
-
In-Ceram
15
-
-
15
-
IPS Empress2
15
-
-
-
-
Tab. 13: Overview on the number of fabricated specimens for load bearing and reliability analyses. All veneered samples had 6.9 mm2 distal interdental connectors.
The framework test setup was designed to allow a minimal mobility of the abutment teeth which simulated the natural teeth mobility. Its schematic drawing is shown in
fig. 94. The major design characteristics include the mobility of posts, and the load application perpendicular to the flat occlusal surface. Posts were cylinders of 7.0 mm and 8.0 mm diameter with a 1.0 mm circular shoulder, a ball end at the bottom side and a conical 6° tapering which exactly fitted to the test frameworks. They were made of hardened steel to minimize their residual deformation due to the test load and were surrounded by a 0.75 mm layer of plastic hose which assured their lateral mobility. The holder of the test setup was made of an aluminium alloy having cylindrical holes of 7.6 mm and 8.6 mm diameter and 16.1 mm hole distance. The framework was always placed unce-
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mented on the abutment posts and loaded with a flat ended rod until fracture. A teflon disk of 5.0 mm diameter and 2.0 mm height, was placed between the rod and the framework to prevent any contact damage and to provide a homogeneous load distribution on the pontic. The tests were made at a cross head speed of 0.5 mm/min on a universal testing machine1 recording the load-displacement data. The fracture load was determined using an electronic maximum load indicator. For fractographic analysis the fractured frameworks were sputtered2 with gold in argon atmosphere for 2.5 min at a pressure of 5 Pa before taking SEM3 pictures.
load
teflon disk framework post Image: load_bearing_of_bridges/ plastic hose test_setup.wmf holder Design: load_bearing_of_bridges/ test_setup.ppt
Image: load_bearing_of_bridges/ Bridge_test_setup.jpg
Fig. 94: Framework test setup in diagram (left side) and in reality (right side). The idea is to simulate failure with a fracture similar to the clinical observed fracture at the interdental connector.
The Weibull statistics [101, 102] were calculated for all series, except for the TZ3YB sets with the 0.5 mm and the 1.0 mm veneer thickness. The load at 63 % failure probability is a measure for the characteristic load bearing capacity F0 of a test set. The reliability is defined as the slope m of the Weibull regression line which is an indicator for the variation in load bearing capacity within a set.
1. 2. 3.
Schenk-Trebel, Ratingen, Germany Balzers SCD 40, Bal-Tec, Balzers, Principality of Liechtenstein Jeol JSM6400, 20 kV, WD 15 mm.
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The FEA of the test setup enables the simulation of its load bearing capacity, mechanical stresses and its deformations. It uses the software package Ansys 5.31. Müller et al. describes the basic ideas of FEA and their application with the software Ansys [103]. The schematic FE model is shown in fig. 95 with the boundary conditions that we used (left side). Its affiliated FEA shape represention was produced by digitizing the cavital and occlusal side of the base model with the 6.9 mm 2 connector (fig. 95, right side). The triangulation of each point cloud lead to a patched surface description of either the occlusal or the cavital frameworks’ side which afterwards were combined to form a shell. A network of tetrahedral finite elements was generated, and the number of elements was raised to 13’000 until the tensile principal stress approximated its maximum. The posts and plastic hoses were modelled within the FEA software. The procedure is described in [104]. The FEA boundary conditions describe the loads and the displacements that we assumed ( fig. 95, left side). The z-displacement was blocked at the tip of the posts and their deformation due to Hertzian stress was neglected although a slight oblateness at the tips was observed. At the outer surface of the plastic hoses we blocked all degrees of freedom because the holder was not moved and no deformation was assumed at the inner surface of the holes. An equal-deformation boundary condition was set for all interfaces which were the framework-to-post, post-to-plastic hose, and teflon disk-to-framework one, respectively. The materials data is compiled in tab. 14, the steel, plastic and the teflon were reasonably estimated, the data of TZP and In-Ceram were taken from the manufacture’s specifications. All materials were assumed to be ideally elastic. In the case of the teflon plastic deformation is neglected due to its minor importance. The framework was loaded homogeneously on the pontic with a pressure of 1 MPa on the cylindrical elasticity of 7.0 mm diameter. The pressure is equal to a net load of 38.5 N. An upscaling factor of 33.25 which corresponds to the ratio of characteristic load (F0 = 1280 N) and applied load (38.5 N) was used to calculate the real stresses. We chose the
maximum tensile principal stress failure criterion as appropriate because ceramics are especially susceptible to tensile stress. Further details of the FE analysis are described in [104].
1.
ANSYS, Inc., Canonsburg, PA 15317, USA, Internet: www.ansys.com
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For In-Ceram the same FE model was used but the materials’ properties were substituted by In-Ceram’s (see tab. 14) and the load was set to 510 N (In-Ceram’s characteristic load).
z y
x
elastic
p
elastic Image: load_bearing_of_bridges/ FEA_model.wmf Design: load_bearing_of_bridges/ B B B test_setup.ppt
B
Image: load_bearing_of_bridges/ FEA_model_color.jpg
A
A
Fig. 95: FEA for the framework test setup. Left side: Mechanical model for analysis with the boundary conditions A: displacement in z direction is blocked and no contact deformation due to Hertzian stress is allowed; B: all displacements in each direction is blocked. Materials data according to tab. 14. Right side: FEA representation with the framework shell consisting of 13’000 tetrahedral elements.
material
elastic modulus [GPa]
Poisson ratio [-]
TZP
210
0.25
In-Ceram
320
0.25
steel
210
0.25
plastic
1
0.25
teflon
1
0.25
Tab. 14: Materials data as used for FEA.
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IV.4.3
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Results
The principal stress distribution in the loaded framework is shown in fig. 96 and
fig. 97 [104-106]. In its mesial-distal cross section it possessed the maximum tensile stress of 340 MPa in two small surfacial spots of the connectors at their gingival side ( fig. 96). The arrows point to the places of maximum tensile stresses exceeding 200 MPa. A concave curved line through the pontic shows equal magnitude of tensile stress in between 53 MPa and 100 MPa. The cavital view shows the maximum tensile principal stresses of 340 MPa at the interdental connectors at the transition to the pontic (fig. 97). Most of the surface in cavital view exhibits a tensile stress between 7 MPa and 53 MPa. For In-Ceram we calculated an equal stress distribution as was found in the case of the TZP. However the maximum tensile stress at the gingival side of the connectors was 133 MPa, and therefore much lower than that in the case of TZP [105, 106].
(a)
(b)
(c) Image: load_bearing_of_bridges\FEA_stress_01_photoshop.gif Design: load_bearing_of_bridges\FEA_stress_01_photoshop.psd
Fig. 96: Principal stress distribution in the loaded framework: (a) mesial - distal crosssection, (b) mesial side of the cross-section, and (c) the distal side of the cross-section. The black arrows indicate areas of maximum tensile principal stress of 340 MPa at 1280 N load.
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(a)
(b)
200
Image: load_bearing_of_bridges\FEA_stress_02_photoshop.jpg Design: load_bearing_of_bridges\FEA_stress_02_photoshop.psd
(c)
Fig. 97: Principal stress distribution in the loaded framework: (a) the gingival side of the framework, (b) the distal connector from gingival side, and (c) the mesial connector from the gingival side. The black and the white arrows indicate areas of maximum tensile principal stress of 340 MPa at 1280 N load.
The minimum, the maximum and the average load bearing capacity of the different test sets are compiled in fig. 98 [107]. The highest average load bearing capacity had the TZ-3YB frameworks with the thickest 10.6 mm2 connectors (1669 N) followed by those with the 6.9 mm2 connectors (1192 N), and then those of the TZ3YB-E (1117 N). The frameworks made of TZ-3YB and TZ-3YB-E showed a superior average load bearing capacity compared to that of IPS Empress2 (558 N) and that of In-Ceram (453 N), respectively.
1061 672 411
673
558
255
TZ-3YB (6.9 mm2, veneered)
In-Ceram (6.9 mm2)
TZ-3YB-E (6.9 mm2)
TZ-3YB (6.9 mm2)
0
TZ-3YB (10.6 mm2)
195
500
Empress2 (6.9 mm2)
453
Image: load_bearing_of_bridges\minimax_load_bearing_capacity.WMF Design: load_bearing_of_bridges\minimax_load_bearing_capacity.ppt Kaleidagraph Data: load_bearing_of_bridges\minimax_load_bearing_capacity.qpc
In-Ceram (6.9 mm2, veneered)
1288 1005
1106
919
1109
1117
1669
1778
Minimum Maximum Average
753
1000
1451 1192
2000 1500
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2294
2500
921
Load Bearing Capacity [N]
201
Fig. 98: Minimum, maximum and average load for the test sets. Each set contains 15 specimens. The horizontal line indicates the assumed 880 N mastication load.
Frameworks made of different materials with a similar shape using the 6.9 mm2 connector design were compared using Weibull statistics (see fig. 99) [105-111]. TZ3YB (F0= 1278 N, m= 8.5) and TZ-3YB-E (F0= 1232 N, m= 4.4) samples fabricated by the DCM process exhibited a superior load bearing capacity and a superior Weibull modulus than that of In-Ceram ( F0= 514 N, m= 2.7) and IPS Empress2 ( F0= 621 N,
m= 3.0). Characteristic load bearing capacity of TZ-3YB and TZ-3YB-E were found to be equal whereas the Weibull modulus of TZ-3YB-E was significantly lower than that of TZ-3YB. The Weibull parameter are compiled in tab. 15.
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lnln(1/(1-P))
2
200
100
1
500
1000
TZ-3YB TZ-3YB-E In-Ceram Alumina IPS Empress2
2000 3000 1.0 0.99 0.95 0.90
0
0.63 0.50
-1Image: load_bearing_of_bridges\different_materials.WMF
0.30 0.20
Design: load_bearing_of_bridges\weibull.ppt Kaleidagraph Data: load_bearing_of_bridges\different_materials.QPC
-2
0.10
-3 -4 4.5
Failure Probability [-]
Load Bearing Capacity [N]
0.05 0.02 5
5.5
6
6.5
7
7.5
8
ln (Load Bearing Capacity) Fig. 99: Load bearing capacity and reliability as a function of the frameworks’ material. Similar shape with a 6.9 mm2 connector cross-section area.
TZ-3YB - frameworks of different connector areas are compared using Weibull statistics (see fig. 100). Those with the larger 10.6 mm2 connector showed superior characteristic load bearing capacity (F0= 1795 N) in relation to the smaller 6.9 mm2 connector (F0= 1278 N). The additional 3.7 mm2 connector area shifted the load bearing capacity by 500 N towards higher values. Despite its lower load bearing capacity the specimens with the smaller 6.9 mm2 connectors exhibited the higher Weibull modulus (m= 8.5) than those with the 10.6 mm2 connectors (m= 6.1). For a list of the Weibull parameters refer to tab. 15.
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lnln(1/(1-P))
2
200
100
500
1000
TZ-3YB (10.6 mm2 ) TZ-3YB ( 6.9 mm2 )
1
2000 3000 1.0 0.99 0.95 0.90 0.63 0.50
0 -1
Failure Probability [-]
Load Bearing Capacity [N]
0.30 0.20 Image: load_bearing_of_bridges\different_connector_areas.WMF -2 Design: load_bearing_of_bridges\weibull.ppt 0.10 Kaleidagraph Data: load_bearing_of_bridges\different_connector_areas.QPC -3
0.05
-4 4.5
0.02 5
5.5
6
6.5
7
7.5
8
ln (Load Bearing Capacity) Fig. 100: Load bearing capacity and reliability as a function of the TZ-3YB frameworks’ connector cross-section area [105, 110].
The effect of veneering the TZP framework on the load bearing is shown in fig.
101. TZP frameworks with different veneering thickness at the gingival side of the interdental connector exhibited only a minor increase in load bearing when compared to the unveneered samples. The mechanically weak porcelain on the TZP framework did not decrease its load bearing capacity. This is not in contradiction to chapter IV.2 as the failure criteria of both analysis’ are different. Furthermore, load bearing was independent of the different veneer thickness.
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2
200
100
500
1000
2000 3000 1.0 0.99 0.95 0.90
lnln(1/(1-P))
TZ-3YB framework TZ-3YB, 0.5 mm veneer 1 TZ-3YB, 1.0 mm veneer TZ-3YB, 1.5 mm veneer Image: load_bearing_of_bridges\veneer_different_thicknesss.WMF 0.63 Design:0load_bearing_of_bridges\weibull.ppt 0.50 Kaleidagraph Data: load_bearing_of_bridges\veneer_different_thickness.QPC
0.30 0.20
-1 -2
0.10 0.05
-3 -4 4.5
Failure Probability [-]
Load Bearing Capacity [N]
0.02 5
5.5
6
6.5
7
7.5
8
ln (Load Bearing Capacity) Fig. 101: Load bearing capacity and reliability as a function of the veneer thickness at the the connector’s gingival side [105].
The load bearing capacity of veneered test bridges made of differnent material is shown in fig. 102 . The TZ-3YB bridges exhibited a load bearing capacity of
F0= 1080.5 N and a Weibull modulus m of 6.1, whereas the In-Ceram samples possessed a F0 of 748.4 N and a m of 3.56. The TZP bridges exhibited a 332.1 N (44.4 %) higher characteristic load bearing capacity and simultaneously a 2.53 (71.1 %) higher Weibull modulus than the In-Ceram specimen.
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lnln(1/(1-P))
2 1
200
100
500
1000
TZ-3YB test bridges In-Ceram Alumina test bridges
2000 3000 1.0 0.99 0.95 0.90
0
0.63 0.50
-1
0.30 0.20
-2
0.10
-3
0.05
Failure Probability [-]
Load Bearing Capacity [N]
0.02 Image: load_bearing_of_bridges\veneered_TZP_and_InCeram.WMF -4 Design: load_bearing_of_bridges\weibull.ppt 4.5 5 5.5 6 6.5 7 7.5 8 Kaleidagraph Data: load_bearing_of_bridges\pure_veneered_TZP_and_InCeram.qpc
ln (Load Bearing Capacity)
Fig. 102: Load bearing capacity and reliability of TZ-3YB and In-Ceram test bridges in comparison. TZ-3YB frameworks were veneered using the experimental W35/11 (1.5 mm thickness), the In-Ceram samples with Vitadur Alpha (1.0 mm thickness).
material
6.9 mm2 connector
10.6 mm2 connector
TZ-3YB
m= 8.54 F0 = 1278.4 N
m= 6.05 F0 = 1794.6 N
TZ-3YB-E
m= 4.38 F0 = 1232.1 N
In-Ceram
0.5 mm veneer
1.0 mm veneer
1.5 mm veneer
-
-
m= 6.09 F0= 1080.5 N
-
-
-
-
m= 2.72 F0= 514.2 N
-
-
m= 3.56 F0= 748.4 N
-
m= 3.02 IPS Empress2 F0= 620.5 N
-
-
-
-
Tab. 15: Weibull parameters of the framework specimens. Characteristic load bearing capacity F0 and reliability (Weibull modulus m).
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The load-displacement behaviour of the TZP frameworks is shown in fig. 103. The load increased continuously with displacement until a brittle fracture of the framwork occured. The high amount of displacement of 1.65 mm until the fractured and its non-linearity were mainly due to the deformation of the teflon disk. For In-Ceram and IPS Empress2 frameworks similar load displacement behaviour were observed.
Load [N]
2000 1500 1000 500
Image: load_bearing_of_bridges\load_displacement_E83.WMF Design: load_bearing_of_bridges\load_displacement_E83.ppt
0 0
0.5
1
1.5
2
Displacement [mm] Fig. 103: Load-displacement diagram of TZP frameworks.
However, the load-displacement behaviour of veneered bridges shows specialities in comparison to that of pure frameworks. Load-displacement diagrams of test bridges are shown in fig. 104 [105]. In-Ceram bridges showed a smooth, continuous increase of load with displacement until the brittle fracture of the whole sample occured. In this case the crack propagated without any stop or diversion through the porcelain and the framework. However, in the case of TZP load relieve peaks were observed with displacement, e.g. the load did not increase continuously and smoothly. We attributed the load relief peaks to a crack "stop and go" mechanism: the crack starts in the porcelain and propagates towards the porcelain-framework interface. There it stops and diverts in lingual and in buccal direction, and then propagates in steps with increasing load on both con-
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nector sides towards the occlusal surface of the bridge. No further damage of the porcelain veneer, like delamination was observed until the fracture of the framework. Hence the TZP bridges show a beneficial "stop and go" failure mechanism where the number of load relief peaks varied for each bridge but was correlated with crack propagation.
Porcelain
Porcelain
In-Ceram
TZP
Image: load_bearing_of_bridges\failure_models.wmfCrack Propagation Failure of Test Bridges along the Porcelain Design: load_bearing_of_bridges\bridge_failure_schematics.ppt with In-Ceram-Framework TZP - Interface
Load
Load
Failure of TZP Frameworks
Crack Initiation in the Porcelain
Displacement
Displacement
Fig. 104: Schematic diagrams of load-displacement behaviour of TZP and In-Ceram test bridges. Right: TZP bridges show a first load drop with increasing displacement due to a crack initiation in the porcelain, and a second one due to crack propagation. Left: In-Ceram test bridges fail catastrophically with no obvious stress relief peaks [105].
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The fracture mode of a TZP framework and a TZP bridge are shown in fig. 105. All fractures happened perpendicularly to the mesial-distal axis of the framework in the interdental connector with the fracture surface curved towards the pontic. Curvature was different for each framework. The TZP bridge failed similarly to the TZP frameworks. This described fracture characteristicwas also observed for all other analysed materials (In-Ceram, IPS Empress2).
Image: load_bearing_of_bridges\ failure_pure_framework.jpg
Image: load_bearing_of_bridges\ failure_veneered_framework.jpg
Fig. 105: Typical fracture of the unveneered and veneered TZP frameworks. Left: Fractured TZP framework. Right: Fractured veneered TZP framework. The fracture always initiated at the gingival side of the connector and propagated towards the occlusal surface of the framework.
The fracture surfaces are shown in fig. 106 for frameworks and in fig. 107 for bridges. TZP possessed a homogeneous fracture surface. In-Ceram and IPS Empress2 showed inhomogenities in the fracture surface, however the failure always originated at the gingival side of the interdental connector. In-Ceram’s failure origin could not be exactly located due to its bi-phase structure. In the case of IPS Empress2 some pores were discovered in the region of the probable initiation of the fracture origin, all of these were prone of to be the origin.
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Image: load_bearing_of_bridges\bridge_failure_surface\failure_surface_of_frameworks_1.jpg Design: load_bearing_of_bridges\bridge_failure_surface\failure_surface_of_frameworks_1.psd Design: load_bearing_of_bridges\bridge_failure_surface\failure_surface_of_frameworks_1.ai
Fig. 106: Fracture surfaces of TZP, In-Ceram, and IPS Empress2 frameworks. Left column: The rectangles indicate the probable region of failure origin within the connector. Right column: Zoom-in of the marked region, the arrow marks the failure origin in the case of TZP. Top row: The TZP framework failed from the connector surface (arrow). Middle row: The In-Ceram framework failed from the gingival side of the connector. The failure origin was difficult to find due to its bi-phased microstructure. Bottom row: The IPS Empress2 framework failed from the gingival side. Several pores (arrows) were prone to be the failure origin.
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For test bridges the failure origin was found at the gingival side of the connector [105]. SEM pictures of the fracture surfaces are shown in fig. 107. The TZP bridges failed finally from its internal TZP-porcelain interface because the typical Hackle region was detected only in the TZP framework but was not found in the veneer porcelain. In contrast the In-Ceram test bridges failed from its gingival porcelain surface, the typical mirror region and the Hackle region are observed in the porcelain layer. They proceeded traversing the interface without any diversion into the In-Ceram framework which indicates that there was no energy dissipated. This supports the observation that there was no "stop and go" mechanism in the case of In-Ceram.
Image: load_bearing_of_bridges\bridge_failure_surface\failure_surface_of_bridges.jpg Design: load_bearing_of_bridges\bridge_failure_surface\failure_surface_of_bridges.psd Design: load_bearing_of_bridges\bridge_failure_surface\failure_surface_of_bridges.ai
Fig. 107: Fracture surfaces of TZP and In-Ceram test bridges (veneered frameworks). Left column: The rectangles indicate the probable region of failure origin within the connector. Right column: Magnification of the marked region with arrows indicating the failure origin. Top row: TZP test bridge failed from its gingival side of the framework connector. Bottom row: In-Ceram test bridge failed from the gingival side of the porcelain.
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Discussion and Summary
The FE analysis shows maximum principal tensile stresses at the gingival sides of the connectors (fig. 96) exactly at the place where the failure origin of the frameworks (fig. 106) was located. Therfore the FE model represents qualitatively the real situation in the framework test setup. The fracture surface seems to be perpendicular to the bands of equal stress that we found in our FE analysis. Kelly et. al. also showed by FE analysis the maximum principal tensile stresses being concentrated in the connector when mobile posts are used [99]. This corresponds exactly to our results. However, the absolute maximum stress value in the TZP test bridges of 340 MPa is more than two times less than the bend strength of TZP. Therefore two conclusions may be possible: (1) the stress value calculated by FE is correct, and the milling in the presintered state lowers the bend strength by introducing flaws in the surface or a high security towards measured TZP’s bend strength applies, (2) the FE-calculated stress is too low. For a three-unit In-Ceram front teeth bridge with box preparation which was loaded by 250 N per-bridge-unit from 60° from palatinel direction, Pospiech et. al. found a 534 MPa maximum stress in the connector (height 3 mm, area 8.8 mm2) [112]. Despite the larger connector area and the lower applied load on the pontic, they found stresses four times higher than we found (133 MPa). We conclude that our FE calculated stresses are too low. The load bearing capacity, strength and reliability of frameworks and test bridges were analysed for different materials. Frameworks of TZP possess a more than two times higher average load bearing capacity than equally shaped In-Ceram and IPS Empress2 frameworks. Test bridges of TZP were found to have a 1.5 time higher load bearing than the In-Ceram Alumina samples. We found a higher strength in combination with a higher reliability for TZP bridges and frameworks compared to those of In-Ceram Alumina and those of IPS Empress2. The combination of both characteristics proves the ability to fabricate all-ceramic bridges with a low variation in strength (in fig. 108) and, therefore, offers more security to the dental-technician to guarantee for his products. The load bearing capacity of test bridges is superior to that of the test bars (chapter IV.2), although the bars had a larger 8 mm2 “connector”-cross-section. A probable reason is that maximum bend momentum resides below the pontic in case of bridges whereas at the
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weak interdental connectors only a smaller momentum applies. In the case of test bars the fracture occurs were the maximum bend momentum resides. Reliability of the bridges was only a little smaller than that of the bars with the TZP in compression layer. Retrospectively, the bridges showed the higher load bearing capacity, and a reliability of the same magnitude. Therefore all statements derived form the bars for static load and reliability were on the safe side and an extrapolation to bridges would have been possible.
10 framework test bridge
Reliability [-]
8
TZ-3YB
TZ-3YB
6
TZ-3YB (10.6 mm2)
TZ-3YB-E
Image: 4 load_bearing_of_bridges\weibull_parameter.wmf In-Ceram Empress2 Design: load_bearing_of_bridges\weibull.ppt In-Ceram
2 0
0
500
1000
1500
2000
Load Bearing Capacity [N] Fig. 108: Weibull parameters of the test sets. All frameworks and test bridges were fabricated with 6.9 mm2 connector cross-section area except one framework set with 10.6 mm2 connector (indicated).
However, considering reality, the intraoral cavity offers one of the hardest hostility for materials. Cyclic loads from varying directions, water (saliva), and wide pH and temperature shifts probably characterize the oral environment in general. Chewing forces are reported being in the range of a few Newtons up to 1000 N [113]. Kelly reports average load from 40 N to 250 N during normal chewing and 500 N to 880 N associated with parafunctional habits (bruxism) and dynamic loading with 800 to 1400 chewing cycles per day [62, 114]. Assuming a minimum 14.7 years’ expected lifetime1 7.5 million
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load cycles are calculated. However, this work assumed preclinical testing with a required load bearing capacity for high bite load of 880 N which is an upper limit, to be on the safe side. Assuming mastication load of 500 N test bridges with In-Ceram Alumina framework show a 60 % failure probability, with IPS Empress2 framework a 40 % failure probability and with TZP framework zero failure probability. Highest mastication loads of 880 N cause 99 % of In-Ceram Alumina frameworks to fail and 94 % of IPS Empress2 frameworks. The TZP frameworks as tested here withstand these highest mastication loads and exhibit only a 4 % failure probability. However, the preclinical tests were performed with static and dry conditions, and neglected dynamic loading, water or saliva presence, pH shifts and temperature influence. Dynamic loading leads to a strong decrease of the bend strength - 70 % reduction in the case of In-Ceram Alumina and up to 50 % for TZP compared with static values ([62, 114, 116-121], principles and overview in prosthodontic application [122]). The comparison with the 55 % bend strength reduction in the case of metal-porcelain [123] is interesting. These benchmarkings derive a lower susceptibility of TZP against dynamic loading than In-Ceram and even metal-porcelain. Nevertheless, preclinical benchmark tests for TZP approaching the intraoral reality more closely are not known. The design of bridges especially concerns the highly stressed interdental connectors. In the case of non-precious alloys Schwickerath recommends connector cross-sections of 10 mm 2 (2.5 mm height, 4.0 mm width) for a 10 mm retainer distance and of 16 mm2 (4.0 mm height, 4.0 mm width) for 30 mm retainer distance, respectively [124]. For IPS Empress2 bridges Sorensen [125, 126] recommends a 20 mm2 (5.0 mm height, 4.0 mm width) connector area between premolars; for In-Ceram Alumina bridges Kappert et. al. [100] and Pröbster [127] recommend 16 mm2 connector area (4.0 mm height, 4.0 mm width). From their FEA results Pospiech et. al. advise smooth and rounded contours of the connector with a 4 mm minimal occlusogingival height [112]. Such oversized connectors using a lot of excessive material limit the clinical indication and lead to a non-satisfactory aesthetic appearance. A cross-section area of 7 mm2 to 10 mm2 may be appropriate for posterior TZP restorations when we consider the experimental load bearing capacity, the high load assumptions of 880 N, the failure probability, and the knowledge about bend strength reduction in dynamic loading mode. This connector de1.
based on a survey of 571 dentists for gold -ceramic restorations (according to Anusavice [115])
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sign is comparable to that recommended for porcelain-fused-to-metal multi-unit bridges, however it definitely is less than that for all other known all-ceramic systems. For anterior restorations which bear lower loads connector sizes may be further reduced. These proposed connector sizes seem to be clinically acceptable. Despite, the lean connector size the TZP frameworks offer superior load bearing capacity combined with a superior reliability. Furthermore for designing connectors it is common sense to prefer smooth and rounded contours and to avoid sharp edges and strong separation between the teeth as they may provoke high peak stresses. The failure of all test bridges and frameworks was always initiated at the gingival side of the connector, its fracture surface was curved towards the pontic and was perpendicular to the mesial-distal axis. This failure mode is identical with the clinical failure of multi-unit restorations as reported by Kelly [99, 114] with its origin at the outer porcelain surface or at the interface between porcelain to framework. In contrast crowns fractured from their internal cavity [96-98]. The observed ex vivo failure corresponds to a clinical failure with static loading. A test setup also with mobile retainers was used by Kappert et. al. [100]. In contrast, their test bridges were cemented onto the retainers. However, the produced fracture and the 703 N average load bearing capacity for In-Ceram bridges using connectors of 4 mm diameter (12.6 mm2 area) is similar to the findings in this work. Pospiech et. al. also found no effect of cementation on average load bearing [128]. Therefore, introduc-
ing the simplifications of avoiding cementation and flattening occlusal surface have a minor or no effect on the load bearing capacity. These assumptions are also supported by our FE results. Moreover, no framework or test bridge failed due to contact or Hertzian pressure [114, 129]. Therefore, the presented results exhibit clinical relevance. In contrast to this, a test setup with fixed retainers generally leads to higher load bearing capacity which overestimates the potential of the material and therefore should be critically analysed [100, 128, 130]. The displacement-load and the failure behaviour of test bridges using a TZP framework were different from those with an In-Ceram framework. In the case of TZP a crack-"stop and go" mechanism in the porcelain was observed. Such a mechanism is beneficial for its clinical application as it announces a prospective failure at an early stage. Crack propagation leads to a load drop in the porcelain and the failure of the test
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bridges is only governed by the TZP framework. This was also confirmed by the fracture analysis. In contrast to the above, In-Ceram test bridges exhibit a brittle failure without a crack "stop and go" mechanism at the interface porcelain to framework. With this failure behaviour, an initial crack in the porcelain initiated in the weak porcelain veneering propagates through the interface and, therefore, the bridges’ failure behaviour is only in part determined by the strong framework but also co-determined by the weak porcelain veneer. Finally, we conclude that the DCM process in combination with TZP materials is suitable for the fabrication of all-ceramic restorations and for application in dentistry. TZP possesses superior materials characteristics such as bend strength and toughness compared to other dental ceramics. Frameworks and test bridges fabricated using the DCM process show a superior strength in combination with a superior reliability in comparison to the In-Ceram and the IPS Empress2 system. The load bearing of TZP restorations exceeds the mastication forces in the posterior region even when using a lean and clinically appropriate connector design. Therefore, preclinical testing results suggests the recommendation of TZP bridges for posterior use.
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IV.5
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Clinical Performance of DCM fabricated Dental Restorations [131]
Clinical studies using all-ceramic dental restorations which were fabricated by the DCM process at the Swiss Federal Institute of Technology, Zurich, Switzerland were performed at the University Zurich, Clinic of Fixed and Removable Prosthodontics and
Dental Materials Science in Zurich, Switzerland. This chapter summarizes the results of the clinical application of DCM fabricated dental restorations and is based on the phdthesis of Sturzenegger [131].
IV.5.1
Introduction
All-ceramic restorations available on the market exhibit higher failure rates than PFM restorations during clinical long term studies [132, 133]. Premolar and molar allceramic restorations were reported to fail less than front tooth restoration [115, 134136]. Although In-Ceram alumina clinically succeeded in its early beginnings [100, 127, 137, 138] the premolar and molar bridges proved a high failure rate after the first year in service [99, 139]. Hence the interdental connectors were oversized in a clinically and an aesthetically unacceptable way to be not less than 16 mm2 better 20 mm2 [125, 126]. Despite this all-ceramic restorations are still only recommended for the anterior region, up to now. Highly translucent all-ceramic restorations made of Dicor, Optec, Ceraperl or IPS Empress produce best aesthetic results but they exhibit weak mechanical properties, and therefore they are only recommended for single crowns in the anterior region [115, 140]. All-ceramic restorations built-up on an opaque framework material such as Hi-Ceram and In-Ceram possesses better mechanical properties but they can not satisfy highest aesthetical requirements. TZP has superior mechanical properties which enables its use
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in the whole denture, it benefits form its ivory-white colour, its ability to be coloured, and its translucency making it desirable in the visible region of the denture. Therefore, the idea evolved to extend the clinical use of TZP to all-ceramic restorations. These restorations may have the potential for application in the complete denture. The goal of this chapter is to prove the clinical feasibility of DCM fabricated bridges in the highly loaded molar region of the denture. Therefore a controlled clinical study was performed in collaboration with the University of Zurich, Clinic of Fixed and Removable Prosthodontics and Dental Materials Science in Zurich, Switzerland. Further potential of the DCM system restorations was investigated by treatment of teeth in all other regions. As a result of the clinical application we established an appropriate preparation for the retainers, the clinical feasibility within the observed time frame, and the clinical indication bandwidth for different types of TZP restorations.
IV.5.2
Material and Method
The patients for the clinical study were selected in order to have a representative section of the swiss population and from the clinical reality in a swiss dental practice. No restrictions on social level, profession, age, sex, or oral hygiene were made. The treatment was carried out by one private dentist, one clinical dentist of the University of Zurich, and some students of dentistry at the University Zurich. Different dental techni-
cians who had professional working knowledge but not necessarily experience with allceramic restorations were responsible for the fabrication of the models and for the coating of the TZP frameworks. Two persons with no experience in the dental field did the machining and sintering of the TZP-frameworks. The treatment was similar to routine in a dental private practice, and it involved two to three appointments from the starting to the final session. The complete clinical step-by-step insertion procedure in an overview is shown in fig. 109. The veneering was performed by the dental technician dealing with the circumstance that we were limited in the colours, and therefore also in the aesthetic results. A standard preparation for nonremovable prosthodontic restorations with ceramic cervical margins according to the accepted rules and technology was chosen to keep the conventional dental treatment plan.
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Conventional treatment of the adjacent teeth, the retainers, and the rest of the denture was also made according to the rules and technology known to the dentist in the case of PFM restorations. Desensitizer was used to close the dentin canals of the prepared vital retainer tooth. No additional special treatment was necessary.
image: clinical_performance/clincial_steps.jpg
Fig. 109: Complete step-by-step clinical insertion procedure of an all-ceramic three-unit bridge in the upper left jaw of the patient (Pictures Courtesy of University of Zurich). The tooth 26 is missing in the initial situation, bridge from 25 to 27.
The tooth preparation was made according to fig. 110 using minimal chamfer radii of 0.65 mm except for the preparation margin, a 1.1 mm axial reduction and a conical preparation with a tapering angle of 6-8 ° for molars and of 10 ° for front teeth, respectively. A clear and simple margin with a shoulder or a chamfer margin design was prepared, tangential margins were avoided. The preparation required a 1.9 mm to 2.0 mm
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occlusal reduction (see fig. 111). As tools we used conventional diamond burs. Afterwards desensitizer was applied to close the dentin canals of the vital retainers. The situation in the patient’s mouth was moulded using polyether material, and then they got provisional restorations during the time needed to fabricate the final restorations.
(a)
(b) buccal
Radius: > 0.65 mm
ca
6°- 8°
90°
1 1.
90°
m
ca 1.1 mm
m
buccal
lingual palatel
90°
Radius: > 0.65 mm
90° ca
10°
Fig. 110: Tooth preparation guidelines: (a) for molars, (b) for front teeth.
veneer 1.9-2.0 mm
Design: clinical_performance/ zahnarzt.ppt
retainer
TZP framework
Fig. 111: Occlusal tooth reduction of 1.9 mm to 2.0 mm is required.
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The dental technician fabricated the master models from the negative mould (see
chapter VI.6). Then, the plastic model of the framework was artistically manufactured in the same manner as for wax technique with a minimum of 0.7 mm wall thickness, a minimum of 7 mm2 interdental connector cross-section area and no undercuts. The TZP frameworks were fabricated as described in chapter IV.3. No further "shape optimization" was necessary if the model was designed well (fig. 112).
Do not grind the framework after after sintering to full density!
Do not grind the interdental connectors of the framework after sintering to full density.
Design: clinical_performance/ zahnarzt.ppt
retainer
Fig. 112: Avoid grinding the framework in the dense state for shape optimization.
The coating with the veneer porcelain was done using three enamel and six dentine colours all based on the experimental W35/111, and it was applied in three to four firing steps according to chapter IV.2.2. The design of the porcelain veneer is shown the cross-section view in fig. 113.
1.
Ivoclar, Schaan, Principality of Liechtenstein
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veneer porcelain minimum 1.2 mm
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TZP framework minimum 0.7 mm
porcelain running out at the cervical margin
zahnarzt.ppt retainer
Fig. 113: Minimum thickness of the framework is 0.7 mm and of the veneer porcelain is 1.2 mm (occlusal). Porcelain can run out at the cervical margin.
For the final cementation the temporary restorations were removed and the abutment teeth cleaned. The fit of the bridges were checked especially at the cervical margins. Then their cavities were sandblasted using alumina and cleaned using ethyl alcohol before they were adhesively cemented1. The medical diagnosis at the cementation appointment included: taking x-ray pictures of the abutment teeth before and after the cementation, testing the vitality of the retainers and the adjacent teeth, diagnosis of the parodontics, and taking photographs from occlusal and side views. The medical diagnosis at the recall appointment included: taking x-ray pictures of the teeth, taking photographs from occlusal and side views, testing the vitality of the retainers and the adjacent teeth, probing of periodont of the retainers, the adjacent teeth and the opposing teeth, expertising the framework (intact, fractured) and the veneer (intact, cracked, chipped), probing the cervical margin (non-probing, probing greater than 50 %, probing lower than 50 %), analysis of marginal discoloration and postcariousity (non-carious, carious), and an elucidation of the patient’s satisfaction.
1.
Panavia TC, Kuraray, Osaka, Japan
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For the clinical study a minimum of 20 three-unit bridges in the molar region after being in service for longer than one year were investigated. Additional patients were treated with single crowns, tree-unit to five-unit bridges in the whole denture using the same procedures to gain more experience.
IV.5.3
Results after the first year
For the clinical study a total of 22 bridges in the molar region were inserted: 21 of them being of three-unit type and one being of a four-unit type (tab. 16). Thereof eight bridges were placed in the maxillary arch, and 14 in the mandibular arch, respectively. All inserted restorations were intact at the first recall after a mean of 385 days (minimum 307 days, maximum 488 days) in vivo. No catastrophic failure of the framework nor a chipping or a crack of the veneer porcelain was observed for any of the restorations. The cervical margin of two bridges was more than 50 % soundable, for eleven bridges less than 50 % soundable and for eight bridges not soundable. All probing depths were judged as clinically inconspicious. No discoloration at the cervical margin nor postcariousity were detected. A vitality loss and fistula creation were observed at one premolar. All patients were satisfied with their restorations.
Location
Insertion Date
Recall Date
Time in Service [d]
Framework Status
Veneer Status
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Case
223
1
14-16
24.4.98
28.7.99
453
intact
intact
x
2
15-17
14.5.98
15.9.99
488
intact
intact
x
3
25-27
28.5.98
17.9.99
477
intact
intact
x
4
45-47
28.5.98
17.9.99
477
intact
intact
x
5
45-47
26.6.98
28.7.99
399
intact
intact
x
6
13-15
02.7.98
28.7.99
391
intact
intact
x
7
24-26
02.7.98
28.7.99
391
intact
intact
x
8
14-16
03.7.98
30.7.99
392
intact
intact
x
9
44-46
09.7.98
30.7.99
386
intact
intact
x
10
25-27
16.7.98
17.9.99
432
intact
intact
x
11
45-47
20.7.98
28.7.99
373
intact
intact
x
12
13-15
20.7.98
28.7.99
373
intact
intact
x
13
35-37
20.7.98
28.7.99
373
intact
intact
x
14
35-37
22.7.98
30.7.99
373
intact
intact
x
15
43-46
23.7.98
28.7.99
370
intact
intact
x
16
34-36
23.7.98
28.7.99
370
intact
intact
x
17
45-47
25.8.98
30.7.99
340
intact
intact
18
34-36
03.9.98
28.7.99
328
intact
intact
x
19
43-45
03.9.98
28.7.99
328
intact
intact
x
20
35-37
04.9.98
28.7.99
327
intact
intact
x
21
34-36
16.9.98
28.7.99
315
intact
intact
22
35-37
17.9.98
30.7.99
307
intact
intact
Probing none
< 50 % > 50 %
x
x x
Tab. 16: Overview of the results of the clinical study after the first recall.
One study case of a three-unit bridge, as an example, shows the initial situation, that after the cementation, and that after 15 months in vivo, respectively (see fig. 114). For its time in vivo no visible changes were observed.
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Image: clinical_performance/case_3u_bridge.jpg Design: clinical_performance/case_3u_bridge.psd
Fig. 114: Study case. Top: Initial situation with missing premolar (tooth 15) (Courtesy: University of Zurich). Bottom: Three-unit bridge (from 14 to 16) in vivo after 15 months.
In addition to the study, 10 single crowns, 17 three-unit bridges, 10 four-unit bridges, and 3 five-unit bridges with two or more abutment teeth were inserted maxillary and mandibulary, anteriorally and posteriorally within the complete denture until the end of 1999. A total of 183 units were fabricated including the clinical study. No clinical failure has been observed up to now.
IV.5.4
Discussion and Summary
The preparation of the teeth was seen to be crucial for a good fit. Undercuts, nonchamfered edges, and parallel preparations produced some overworking expense in the range from 5 min to 45 min at the dense sintered framework. This is equal or less compared to that usually needed for cast metal frameworks. The fitting was improved continuously during the study and was judged to be sufficient by the dental technicians. To assure a good fit the dentist should prepare a conical with a 6 ° to 8 ° tapering angle,
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avoid any parallel preparations and any undercuts, use a shoulder or a chamfer margin design, reduce the retainer by 1.0 mm to 1.1 mm in lateral direction and by 1.9 mm to 2.0 mm in occlusal direction and prepare a clean, simple tooth shape and cervical margin line, respectively. The tooth preparation for the dentist is similar to PFM restoration, however he should follow the guidelines to avoid any unnecessary overworking of the framework. The plastic models were also crucial for the fit and require to be net-shaped for the DCM process in order to omit any hard machining step of the sintered framework. This is one difference to other available systems where the dental technician does substantial overworking of the metallic or the ceramic framework. The DCM process requires models of similar shape as commonly used for the lost wax technique in case of PFM restorations. However for the clinical studies the wax was substituted by a plastic material, but optical digitizing technology will permit the used of favoured dental waxes. During the fabrication process of a framework the dental technician performs unliked work such as casting the moulds, casting the metal, demoulding the part and cleaning the metallic framework. Especially these steps were substituted by an automated fabrication of a fitting TZP framework. Nevertheless the remainder of the restoration fabrication process stayed the same for the patient, the dentist and the dental technician. The dentists and the dental technicians note this as a real benefit. The dentist can use the same knowledge and techniques, and the dental technician may focus to the artistic design of restorations. The aesthetic appearance is determined by the framework’s colour and its translucency. For the clinical study the TZP was only available in ivory-white which exhibited a translucency as shown in fig. 115. The veneer porcelain was at hand in six dentine colours and three enamel colours, respectively. Despite these temporary shortcomings we achieved satisfying aestethic results.
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Image: clinical_performance/transluscency.jpg Design: clinical_performance/transluscency.psd
Fig. 115: Translucency of an all-ceramic three-unit bridge (Courtesy of University of Zurich). Right: Veneered bridge in front of an aperture with a shining light bulb behind. Left: Light transmittance through the veneered TZP framework.
In addition to the clinical study, a number of single crowns, four-unit bridges and five-unit bridges were fabricated and inserted to investigate the DCM system’s indica-
tion bandwidth. In the following three clinical cases are reported to demonstrate the possibilities of the DCM system. All restorations were inserted using the same methods as used for the clinical study.
Case 1: A full restoration of the lower jaw is shown in fig. 116, fig. 117 and fig.
118. The patient had abraded his natural tooth substance due to bruxism disease. The front teeth 31, 32, 41, and 42 in the lower jaw were treated using zirconia root posts with a ceramic build-up1 and tooth 33 was treated by a conventional crown, respectively (fig.
116). Five front tooth crowns in the lower jaw which are the smallest teeth in our denture and therefore the most difficult to produce were fabricated using the DCM process.
1.
Ivoclar, Schaan, Principality of Liechtenstein
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Image: clinical_performance/bruxism_1.jpg Design: clinical_performance/bruxism_1.psd
Fig. 116: Crowns for a bruxism patient who had abraded his natural tooth substance (Courtesy of University of Zurich). Upper left: Initial situation in the lower jaw. Upper right: Treatment with zirconia posts. Bottom left: Veneered crowns fabricated using DCM process. Bottom right: Final situation in the lower jaw with the restored teeth (42, 41, 31, 32, and 33).
For the same patient a four-unit and a three-unit bridge were fabricated (see figure
117). We used similar connector dimensions to those for PFM restorations. Although bruxism is a strict contraindication for ceramic restorations this patient was treated mandibularily completely with DCM fabricated TZP restorations. The situation before, directly after cementation and 12 months in vivo is shown in fig. 118 from its facial view. No visible change was observed during 12 months in vivo.
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Image: clinical_performance/bruxism_2.jpg Design: clinical_performance/bruxism_2.psd
Fig. 117: A four-unit (43-46) and a three-unit bridge (34-36) fabricated using the DCM process for a bruxism patient (Courtesy of University of Zurich). Upper left: TZP frameworks for a four-unit bridge (43-46) and a three-unit bridge (34-36) from occlusal view. Upper right: Four-unit framework from vestibular view. Middle left: Veneered four-unit bridge. Middle right: Lower jaw from occlusal view after cementation of the restorations. Bottom left: Final situation of the four-unit bridge from right lateral view. Bottom right: Final situation of the three-unit bridge from left lateral view.
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Image: clinical_performance/bruxism_3.jpg Design: clinical_performance/bruxism_3.psd
Fig. 118: Initial situation (left) of a bruxism patient versus the final situation after insertion with all restorations in place (centre), and the lower jaw 12 months after insertion (right) (Courtesy of University of Zurich).
Case 2: A five-unit bridge from tooth 13 to tooth 17 possessing one mesial retainer, one distal retainer and three-joint pontic units was fabricated using the DCM process. The situation is shown in fig. 119. The large span-width between the retainers make this restoration extraordinary. Furthermore, the antagonistic situation required a sophisticated design of the framework with a curvature in two directions (bucco-palatel and craniocaudal).
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Image: clinical_performance/5u_bridge.jpg Design: clinical_performance/5u_bridge.psd
Fig. 119: Large span-width five-unit bridge possessing two retainers and in between three-joint pontic units from tooth 13 to tooth 17 (Courtesy of University of Zurich).
Case 3: Additional restorations including multiple-unit restorations with more than two retainers were fabricated and inserted. Examples are shown in fig. 120 which prove the possibilities of the DCM process for even more complex shapes and for long spanning restorations.
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Image: clinical_performance/5u_bridge_w_3_retainer.jpg Image: clinical_performance/5u_bridge_w_4_retainer.jpg
Fig. 120: Five-unit bridge with four retainers and one pontic unit in presintered state (left side) and one with three retainers and two pontics in sintered state (right side).
These three cases are successfully treated using the DCM system. All of these restorations were still in service and no failures nor any other problems are known up to now. Other available all-ceramic systems were not able to cover these cases which may prove the superiority of the DCM system compared to the other systems. Considering the three demonstration cases we conclude that the clinical indication for the DCM process may be given for the complete denture, even in case of high loads, of bruxism disease or of long span restorations. One other case (case no. 8, tooth 14) needed endodontical treatment after cementation. This may have been caused by the preparation and may not be linked to the TZP framework. Endodontical treatement requires to make a hole though the restoration into the tooth root. This was performed using high-speed handpieces (Airotor) equipped with a rough-grained diamond bur (80 µm). After the conventional endodontical treatment the hole was closed with a composite material. Post treatment is possible without destroying or removing the whole TZP restoration. Clinical studies concerning the failure rates, the indication bandwidth and the con-
nector design are found in the literature for other available all-ceramic systems. Negative preclinical results with other all-ceramic systems such as Hi-Ceram1, Optec2 and
Dicor3 were reported by Elmiger [141] and Wohlwend [140] and lead to their strict contra indication for the posterior region. All-ceramic crowns fabricated by the Procera4 1.
Vita Zahnfabrik, Bad Säckingen, Germany Jeneric / Pentron, Wallingford, CT, USA 3. Densply DeTrey, Dreieich, Germany 2.
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process produced successful clinical results after five years of service with a 95 % survival rate [142]. If failures occurred they were predominantly for posterior restorations. Scotti reported a successful clinical investigation with a survival rate of 98 % using In-
Ceram crowns, but they were only placed in the anterior region up to the premolars [143]. Pröbster showed a 100 % survival rate for 61 full-coverage In-Ceram crowns placed either in the anterior or the posterior region for a mean period of 20 months (between 4 and 35 months). Furthermore he reported a 93 % survival rate for 15 bridges of the three-unit to five-unit type which were placed either anteriorally or posteriorally within a short-termed observation period of 12 months [127]. His bridges were designed according to Kappert [100] with thick interdental connectors of at least 16 mm2. During 12 months Sorensen [126] investigated 41 three-unit bridges made of IPS Empress2. Despite their large 16 mm2 connectors and the restriction to use the second canine as outer most distal retainer four bridges (10 %1) failed within the first year. To sum up: the clinical studies were mostly performed with the all-ceramic restorations placed in the anterior region. Only a few studies report on the placement in the molar region and if so, clinical failures were found. The demand for the highest load bearing is for bridges in the posterior regions as we investigated here. Despite this all TZP restorations are still in service without failures. Some of the studies mentioned above relied on investigations extended over a longer time period. Having the first bridge placed April 24th, 1998 no problems are known to date. Nevertheless the clinical study is ongoing to gain long term experience as well. Furthermore our clinical results may not be extrapolated and be applied to TZP restorations fabricated using hard machining processes as the DCS2 or the Girrbach Dental3 systems do (see chapter III.6). This type of machining may decrease strength and reliability of fabricated restoration and were not investigated here.
4.
Nobelpharma Biocare, Goteborg, Sweden Sorensen placed a total of 60 bridges from which 41 bridges were of a design similar to that used in this work. We included only those 41 bridges coming to a higher failure rate than him. 2. DCS President, Allschwil, Switzerland 3. Girrbach Dental, Pforzheim, Germany 1.
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This is the first known clinical study for TZP restorations. From its very promising results we conclude that the indication of all-ceramic restorations in the posterior region may be possible. DCM enables the successful clinical application of TZP frameworks for single crowns, three-unit up to five-unit dental bridges with two or more retainers in the complete denture. Even in the cases of bruxism the first years analysis is very promising. Furthermore, the treatment steps for dentists stay the same and the fabrication steps for the dental technician experience only change minorly. All patients wearing DCM restorations were satisfied concerning the aesthetics and the biocompatibility and did not experience the known phenomenon of metallic taste nor that of the hot-cold-hypersensitivity. Our initial question “Do TZP all-ceramic dental bridges survive in the posterior region?” can be answered absolutely positivly after the first year in service. Long-term clinical results are not yet available, are however under investigation.
IV.6 [1]
References
Zheng, J. and Reed, J.S.: "The Different Roles of Forming and Sintering on Densification of Powder Compacts." The American Ceramic Society Bulletin, 1992. 71(9): p. 1410-1416.
[2]
ISO: "Dentistry - Ceramic Denture Teeth.", in ISO 4824. 1993, International Organization for Standardization. p. 1-4.
[3]
Kassenzahnärztliche Bundesvereinigung (KZBV) ( ed. "KZBV Jahrbuch 96. Statistische Basisdaten zur vertragszahnärztlichen Versorgung." Third Edition ed. 1996, Deutscher Ärzte Verlag: Köln.
[4]
Schumacher, G.-H.: "Odontographie: Eine Oberflächenanatomie der Zähne.", 4. ed. 1983, Leipzig: Johann Ambrosius Barth.
[5]
With, G.d.: "Process Control in the Manufacture of Ceramics." in Processing of
Ceramics. Part I. , R.J. Brook, Editor. 1996, VCH Verlagsgesellschaft mbH: Weinheim, New York, Basel, Cambridge, Tokyo. p. 27-67. [6]
Reed, J.S.: "Principles of Ceramics Processing.", 2nd Edition ed. 1995, New york: John wiley & Sons, Inc.
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[7]
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Richerson, D.W.: "Modern Ceramic Engineering. Properties, Processing, and Use in Design.", 2nd Edition, Revised and Expanded ed. 1992, New York: Marcel Dekker, Inc.
[8]
Bortzmeyer, D.: "Die Pressing and Cold Isostatic Pressing." in Processing of Ce-
ramics. Part I., R.J. Brook, Editor. 1996, VCH Verlagsgesellschaft mbH: Weinheim, New York, Basel, Cambridge, Tokyo. p. 127-152. [9]
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Zentrum für Zahn-, Mund- und Kiefernheilkunde der Universität Zürich, Klinik für Kronen- und Brückenprothektik, Teilprothetik und zahnärztliche Materialkunde. Universität Zürich: Zürich. [132] Lehner, C.R., Studer, S., and Schärer, P.: "Full Porcelain Crowns made by IPS Empress: First Clinical Results." Journal of Dental Research, 1992. 71: p. 658. [133] Moffa, J.P.: "Clinical Evaluation of Dental Restorative Materials.", 1988, Letterman Army Institute of Research: San Francisco. [134] Coornaert, J., Adrianes, P., and deBoever, J.: "Longterm Study of PorcelainFused-to-Gold Restorations." Journal of Prosthetic Dentistry, 1984. 51(3): p. 338342. [135] Kerschbaum, T. and Voss, R.: "Guss- und metallkeramische Verblendkrone im Vergleich - Ergebnisse einer Nachuntersuchung bei Teilprothesenträgern." Deut-
sche Zahnärztliche Zeitung, 1977. 32: p. 200-206. [136] Leempoel, P.J.B., Eschen, S., deHaan, A.F.J., and Hof, M.A.V.t.: "An Evaluation of Crowns and Bridges in a General Private Praxis." J Oral Rehabil, 1985. 12: p. 515-528. [137] Kern, M., Knode, H., and Strub, J.R.: "The All-Porcelain, Resin-Bonded Bridge."
Quintessence International, 1991. 22(4): p. 257-262. [138] Sorensen, J.A., Knode, H., and Torres, T.J.: "In-Ceram All-Ceramic Bridge Technology." Quintessence Dent Technol, 1992. 15: p. 41-46. [139] Lüthy, H., Filser, F., Gauckler, L.J., and Schärer, P.: "High Reliable Zirconia Bridges by Direct Ceramic Machining Process (DCM)." Journal of Dental Re-
search, 1998. 77(Special Issue): p. 762. [140] Wohlwend, A., Strub, J.R., and Schärer, P.: "Metal-Ceramic and All-Porcelain Restorations: Current Consideration." International Journal of Prosthodontics, 1989. 2(1): p. 13-26. [141] Elmiger, R., Hagmann, A., Wohlwend, A., and Schärer, P.: "Vollkeramikbrücken: eine Utopie oder ist die Inkorporation solcher Brücken sinnvoll?" Die Zahntech-
nik d. Schweizer Zahntechn. Vereinigung, 1989. 46(15): p. 15-18.
Direct Ceramic Machining Process
246
[142] Oden, A., Andersson, M., Krystek-Ondracek, I., and Magnusson, D.: "A 5-year Clinical Follow-up Study of Procera AllCeram crowns." Journal of Prosthetic
Dentistry, 1998. 80(4): p. 450-455. [143] Scotti, R., Catapano, S., and D'Elia, A.: "A Clinical Evaluation of In-Ceram Crowns." International Journal of Prosthodontics, 1995. 8(4): p. 320-323.
247
V.
Outlook
Outlook
During this work, an experimenal DCM system has been developed, its feasibility has been proven, and the system is now available for the fabrication of dental restorations. However, three directions for further research and development are suggested, concerning a more thorough investigation of certain materials and processing details, the commercialization of the DCM system for dental application, and its extension to other materials and applications. As for the first direction, the details to be further investigated are the homogeneity, the coloration and the color evolution, the machining mechanisms, and the intrinsic material properties. In this work, the homogeneity was determined on a macroscopic scale by density measurements of rings and indirectly by pore size distribution using mecury intrusion porosimetry on a microscopic scale. Local density and its spatial distribution were determined in a non-destructive way by use of computer tomography [1-6], which allows to link blank homogeneity and dimensional tolerance. Continuing on this way may provide know-how for increasing further the accuracy and net-shaping abilities of the DCM process. The coloration of the blanks by cation infiltration has already been principally established. Dental application requires a tooth-like color palette and homogenous color penetration of the framework. Completely penetrating the blank while controling the evolution of color was not possible so far due to lack in understanding of the underlying principles. A clarification could be gained by investigating the cation penetration mechanism into the TZP blanks as a function of kind of cation, its concentration and its penetration time. In addition, optimization of the sintering temperature as a major control mechanism for the evolving color is required.
Outlook
248
Although the machining has been thoroughly investigated in this work, its basic mechanisms require further analysis. We found that material removal changes from cutting to brittle chipping depending on the processing of the blanks. Differences in material removal characteristics of TZP (this work) and of alumina [7, 8] were found. Better insight into machining would provide means to its optimization from the point of view of production engineering and simultaneously from that of materials engineering. This would allow the adjustment of the machinability to other materials as long as similar removal mechanisms apply. The second direction for further development, commercialization of the DCM system, requires several additional steps. A machining and digitizing system dedicated to dental application needs to be developed; the number of patients to be treated needs to be increased; extend the clinical longterm studies should be extended; the DCM system has to be benchmarked against currently available systems; and, approval has to be gained from the European Committee for Standardization (CEN / CENELEC) and from the United States Food and Drug Administration (FDA) to market the system in the respective regions. Commericialization also requires a profound understanding of the aging effects of the intraoral situation on the materials, in particular thermo- and load-cycling, salvia corrosion and pH-effects. The third direction of future development should be towards broadening the pos-
sible application of the DCM process in dentistry and also in other technical fields. In our work, the DCM system was used only for the fabrication of crowns and bridges. However, dentistry would have further benefits from a broader use, such as the machining of supraconstructions which have to be removably mounted to dental implants. Concerning non-dentistry applications, ceramics are increasingly used in smallscaled, complex-shaped products.As an example, the fabrication of an alumina-based high-temperature micro furnace for crystallographic x-ray diffraction investigation [7] showed the benefits of DCM in terms of faster product development at lower costs. However, technical applications often require significantly smaller tolerances than dental applications, and they also use other materials such as alumina, silicon nitrid, silcon carbid or zirconia-toughened-alumina.
249
V.1 [1]
Outlook
References
Kong, C.M. and Lannutti, J.J.: "Localized Densification during the Compaction of Alumina Granules: The Stages I-II Transistion." The American Ceramic Society
Bulletin, 2000. 83(4): p. 685-690. [2]
Lu, P., Lannutti, J.J., Klobes, P., and Meyer, K.: "X-Ray Computed Tomography and Mercury Porosimetry for Evaluation of Density Evolution and Porosity Distribution." Journal of the American Ceramic Society, 2000. 83(3): p. 518-522.
[3]
Lu, P.K. and Lannutti, J.J.: "Effect of Density Gradients on Dimensional Tolerance during Binder Removal." Journal of the American Ceramic Society, 2000. 83(10): p. 2536-2542.
[4]
Lannutti, J.J., Deis, T.A., Kong, C.M., and Phillips, D.H.: "Density Gradient Evolution during Dry Pressing." The American Ceramic Society Bulletin, 1997. 76(1): p. 58.
[5]
Rabe, T., Goebbels, J., and Kunzmann, A.: "Local Resolution Density Determinations on Die Pressed Green Bodies." cfi/Ber. DKG, 1998. 75(6): p. 19-23.
[6]
Phillips, D.H. and Lannutti, J.J.: "Characterization of density gradients during dry pressing of ceramics.", 1993, Ohio Supercomputer Center, Ohio State University, 1224 Kinnear Road, Columbus, OH 43212-1163, USA: Columbus.
[7]
Estermann, M., Reifler, H., Steurer, W., Filser, F., Kocher, P., and Gauckler, L.J.: "A High-Temperature Furnace for X-Ray Diffraction with Directly Machined Alpha-Al2O3 Ceramic Part." Journal of Applied Crystallography, 1999. 32: p. 833836.
[8]
Schippers, C.: "Grünbearbeitung von Oxidkeramik mit definierter Schneide.", PhD, 1999, in Fakultät für Maschinenwesen. Rheinisch Westfählische Technische Hochschule: Aachen.
Outlook
250
251
VI. VI.1
Appendix
Appendix
Chemical analysis of TZP powder lots
Lot-No
Y2O3
Al2O3
SiO2
Fe2O3
Na2O
IgnitionLoss
Z303046B
5.02
< 0.005
0.005
0.002
0.020
3.57
Z304106B
5.09
< 0.005
0.009
0.002
0.018
3.53
Z306283B
5.14
< 0.005
0.006
0.002
0.021
3.48
Z306322B
5.12
< 0.005
0.007
0.002
0.022
3.47
Z307385B
5.13
< 0.005
0.006
0.002
0.022
3.30
Z307386B
5.12
< 0.005
0.006
0.002
0.022
3.25
Z308658B
5.14
< 0.005
0.002
0.002
0.023
3.49
Z309112B
5.10
< 0.005
0.002
0.004
0.015
3.45
average [wt-%]
5.108
< 0.005
0.005
0.002
0.020
3.443
variance [%]
2.35
9.30
Tab. 17: Chemical analysis in wt-% of used lots of TZ-3YB powder (according to Tosoh).
Lot-No
Y2O3
Al2O3
SiO2
Fe2O3
Na2O
IgnitionLoss
Z303046B
5.18
0.241
0.006
0.002
0.023
3.36
Tab. 18: Chemical analysis in wt-% of used lots of TZ-3YB-E powder (according to Tosoh).
Appendix
252
VI.2
Methods
VI.2.1
Density of Green Bodies
The overall average density of the green bodies is determined from its dimensions and its mass. The cylindrically shaped green bodies or blanks were weighed on a precision balance1 and then measured three times in diameter and height using a digital vernier calliper2. In order to get a minimum variation in weight each blank was optically checked for undamaged edges.
VI.2.2
Mercury Intrusion Porosimetry
Compacted powder is filled in a pressure vessel which afterwards is completely filled up with mercury3. External pressure forces the fluid mercury to intrude into the pores. At low pressures only the big pores are filled with mercury, with increasing pressure also the smaller pores are filled. The amount of pores at a specific pore size is proportional to the changement of the intruded mercury volume. Calculation of pore radius assuming cylindrical pores was performed using the Washburn equation [1] e.g. the equilibrium of capillary force (eq. 14) and external pressure on circles (eq. 15). Parameters are: r (pore radius), σ: (surface tension of mercury, 480 mN/m), θ: (contact angle of mercury, 141.3 °), and p (absolute pressure).
1.
capillary force = – 2 ⋅ π ⋅ r ⋅ σ ⋅ cos θ
Eq. 14
2 external pressure force = π ⋅ r ⋅ p
Eq. 15
AE 200 Delta -Range, Mettler Toledo, Greifensee, Switzerland DigiCal, Brown & Sharpe Tesa SA, Renens, Switzerland 3. Porosimeter 2000, Carlo Erba Instruments, Rodano, Italy 2.
253
VI.2.3
Appendix
Hardness
Hardness was determined by means of the indenter method1 using eq. 16 for calculating the Vickers hardness value HV in MPa with load F in Newton and the diagonal length d as average of both diagonals of the pyramid indent in µm2.
6 F HV = 0.189 ⋅ 10 ⋅ ----2 d
VI.2.4
Eq. 16
Fracture Toughness
Fracture toughness is evaluated by the critical stress intensity factor KIc for mode I loading which was determined using the indenter method [3-5]. The KIc values were calculated according to Anstis et. al. using eq. 17 [6, 7], as well as according to Niihara et. al. using eq. 18 [8]. Parameters in the equations are: hardness H in MPa, elastic modulus
E in MPa, a impression half-diagonal length in m, c average length of radial cracks in m, l average length of Palmqvist cracks in m and F load in N, respectively. Hardness H was calculated using eq. 19 which differs from the Vickers hardness HV calculation mentioned. Anstis et. al require c/a values being being larger than c/a ≥ 2 for using eq.
17 which was not fulfilled. Niihara et. al restrict the use of eq. 18 to l/a values being in the range from 0.25 ≤ l/a ≤ 2.5 which was valid for all measured cracks and indents. However both values were calculated.
1.
Zwick type 3212001/00 equipped with diamond pyramid, Zwick, August-Nagel-Strasse 5, 89079 Ulm, Germany 2. Two different definitions of Vickers hardness are common differing by 7.9 % according to Wachtman [2]. One of them uses the area of contact of the four faces of the indenter pyramid, the other uses the projected area of the indenter contact. Here the first definition is used.
Appendix
254
a 1.5 E 0.5 K Ic = 0.032 ⋅ H ⋅ a ⋅ --- ⋅ ---- c H a K Ic = 0.018 ⋅ H ⋅ a ⋅ --- l
0.5
E ⋅ ---- H
Eq. 17
0.4
Eq. 18
F H = ------------2 ⋅ a2
Eq. 19
VI.3
Calculation and Estimation of Errors
VI.3.1
Enlargement Factor absolute measuring magnitude of the error variable
relative error [%]
diameter d
0.01mm
25 mm
0.04
height h
0.01 mm
45 mm
0.02
mass m
0.001 g
70 g
0.003
density water ρw
0.0002 g/cm3
1 g/cm3
0.02
weight (in air) A
0.0002 g
3g
0.007
weight (in water) P
0.0002 g
0.4 g
0.05
Tab. 19: Error assumptions for the variables in eq. 20 to eq. 23
255
Appendix
δf = 1 p + δρ --- ⋅ δρ -------------------s- 3 f ρp ρs
Eq. 20
δρ p = δ d + δ h + δ m ------------------------ρp d h m
Eq. 21
δρ s = δ A + δ P + δρ w ---------------------------A P ρs ρw
Eq. 22
Eq. 23
δf = δls + δlp ----------------f ls lp
relative measurement error [%]
δf ---f
0.05
δρ s --------ρs
0.08
δρ p --------ρp
0.06
Tab. 20: Errors in enlargement factor’s equations.
Appendix
256
VI.3.2
Bend Strength
absolute measuring magnitude of the error variable
relative error [%]
load F
0.5 N
300 N
0.02
height h
0.005 mm
2 mm
0.25
width b
0.005 mm
4 mm
0.13
span l
0.1 mm
15 mm
0.67
bend strength σ0
1.32
Tab. 21: Error assumptions for the variables in eq. 24.
δσ 0 δF δl δb δh --------- = ------ + ---- + ------ + 2 ⋅ -----F l b h σ0
VI.3.3
Eq. 24
Density by Archimedes method
absolute measuring magnitude of the error variable
relative error [%]
water temperature ϑ
1K
density water ρw
0.0002 g/cm3
1 g/cm3
0.02
weight (in air) A
0.0002 g
2g
0.01
weight (in water) P
0.0002 g
0.3 g
0.07
density ρs
0.006 g/cm3
6.05 g/cm3
0.1
Tab. 22: Error assumptions for the variables in eq. 25.
257
δρ s = δ A + δ P + δρ w ---------------------------A P ρs ρw
VI.3.4
Appendix
Eq. 25
Vickers Hardness
The relative error of Vickers hardness is calculated according to eq. 26 assuming the calibrated weight error F as negligible. Error of the length of the indentation diagonal d is estimated to 2 µm. For the green body d is in the range of 300 µm to 500 µm and therefore the relative error is estimated to 1.5 %. For the sintered body d is in the range of 50 µm to 100 µm, depending on the load, and the relative error is estimated to 8 %.
δ HV = 2 ⋅ δ d ---------------HV d
VI.3.5
Eq. 26
Fracture Toughness
The relative error of fracture toughness is calculated according to eq. 27 and eq. 28 assuming the load error as negligible. Absolute error of a, c, l is estimated to 2 µm, the relative error of the elastic modulus is set to 5 %. Absolute and relative errors of the KIc according to Anstis et. al. and to Niihara et. al. are listed in tab. 23.
δ KIc -----------KIc δ KIc -----------KIc
F + 5 ⋅ δ a + 1.5 ⋅ δ c + 0.5 ⋅ δ E = 1.5 ⋅ δ-------------------F a c E Anstis
Eq. 27
F + 3.5 ⋅ δ a + 0.5 ⋅ δ l + 0.4 ⋅ δ E = 1.4 ⋅ δ------------------F a l E Niihara
Eq. 28
δH δF δa ------- = ------ + 2 ⋅ -----H F a
Eq. 29
Appendix
258
Load [N]
KIc [%] KIc[MPa√m] KIc [%] KIc [MPa√m] (Anstis et. al.) (Anstis et. al.) (Niihara et. al.) (Niihara et. al.)
10
-
-
-
-
20
± 24
±2
± 23
±2
30
± 20
±1
± 17
±1
40
± 17
±1
± 14
±1
50
±9
± 0.5
±7
± 0.5
Tab. 23: Absolute and relative errors for the fracture toughness.
VI.4
TZ-3Y microstructure
Image: figures\Beatrix_Michel\michel_01.jpg
Fig. 121: Microstructure of TZ-3Y [9]. Chemical composition is SiO2: < 0.03 wt-%, Al2O3: 0.28 wt-%, Y2O3: 5.36 wt-% and Na2O: 0.08 wt-% is similar to TZ3YB-E, but its mean grain size is larger than TZ-3YB-E by a factor 3.
259
VI.5
Appendix
Transmittance by Bouguer-Lambert Law
The absorption of electromagnetical waves in a homogneous medium is described by the Bouguer-Lambert law [10]. Light is an electromagnetical wave, but the law applies only for the monochromatic light. According to this law the initial intensity I0 is exponentially decreased to I(ω,x) depending on the lightwave frequency ω and on the medium’s thickness x (eq. 30, eq. 31). For the curve fitting we used eq. 32.
I ( ω, x ) = I0 ⋅ exp ( – α ( ω ) ⋅ x )
Eq. 30
4 ⋅ π ⋅ κ(ω) α ( ω ) = --------------------------λ
Eq. 31
x I( x) T ( x ) = --------- = A ⋅ exp – --- I0 B
Eq. 32
with α being the coefficient of absorption in cm-1, κ(ω) the dimensionless materialspecific absorption index , and λ the wave length in cm, respectively. T is the relative transmittance, and A, B are the regression coefficients.
VI.6
Fabrication of plastic models for frameworks
The process of fabrication the plastic framework models by a dental technician is shown in fig. 122. After preparation of the teeth by grinding the dentist made a silicon mould1 (left side of (a)) similar to the procedure as in case of PFM restorations. This silicon mould represents the negative of the situation in the mouth. A positive model of the upper and lower jaws was produced (right side of (a)) by pouring a slurry of super hard
1.
Permadyne, ESPE, Seefeld, Germany
Appendix
260
gypsum 1 into the moulds. Then the region around the prepared abutment teeth was sawed into pieces. The gypsum models of the upper and the lower jaw were mounted in the articulator (b). By adjusting the occlusion the dental technician recognizes the available design space for the veneered bridge and the framework. Undercuts or defects in the gypsum were carefully blocked-out with wax. The preparation margin of each tooth was marked with a fine pencil using a binocular microscope (c) because the cervical regions have to be fabricated very accurately. A gypsum hardener2 was brushed onto the sawed tooth (d), and dried before a separating agent is applied to the gypsum tooth (e). Then the artistic design of the framework model is started using Targis-Base composite3 which has to be polymerized afterwards. After having fabricated the crown copings the pontic was modelled in order to complete the framework model (g). The last step was to check the occlusion of the framework and the design space for the veneer within the articulator (h). No further grinding of the TZP framework should be necessary to optimize its shape. Therefore, the framework fabrication has to be performed carefully and accurately.
1.
Fuji-Rock, GC, Leuven, Belgium Margidur, Benzer, Zürich, Switzerland 3. Ivoclar, Schaan, Principality of Liechtenstein 2.
261
(a)
Appendix
(b)
image: plastic_framework/plastic-framework-2.jpg design: plastic_framework/plastic-framework-2.psd
(c)
(d)
(e)
(f)
(g)
(h)
Fig. 122: Fabrication of plastic framework model for a three-unit bridge. (Courtesy of University of Zurich)
Appendix
262
VI.7
Stress in bilayered structures
VI.7.1
Stress due to mechanical loading
Stress σmax in the porcelain edge fibre of the bilayered structure was calculated using the mathematical model presented in chapter IV.2.2. The results of the calculation are shown in fig. 123. The shape of the curves for both porcelain thicknesses is similar to the calculated bend strength (see fig. 46) except that the absolute stress differences became smaller.
300
W35/9
W35/3
W35/4 W35/11
W35/10
σmax [MPa]
250 200
MOR of pure porcelain (64 MPa)
150
image: chapter_3/Veneer_Porcelain/s_max.wmf design: chapter_3/Veneer_Porcelain/load_bearing.ppt
100 50 0
Porcelain 1 mm Porcelain 3 mm
-2.0
-1.5
-1.0
-0.5
0.0
0.5
1.0
∆ TEC [10-6 K-1] Fig. 123: Stress in the porcelain edge fibre at the first crack in the porcelain. ∆TEC = TECporcelain - TECzirconia.
263
VI.7.2
Appendix
Stress due to TEC mismatch
Residual stress σTEC due to TEC mismatch in the edge fibre of the porcelain layer was calculated according to [11]. The model assumes a linear elastic behaviour (Hook), a pure bend loading, and that the bar’s cross-sections remain planar. The elastic moduli
E of TZP and of porcelain were set to 205 GPa and 70 GPa, respectively and a glass transition temperature Tg of 580 °C was assumed for all porcelain mixtures. The application temperature Ta was set to 37 °C. The residual stress σTEC shows a linear dependence ∆TEC. For TEC of porcelain smaller than the TEC of TZP (∆TEC < 0) compressive residual stresses σTEC were calculated in case of the 1 mm thick porcelain layer whereas residual tensile stress σTEC is present in case of the 3 mm thick the layer. In the case of the 1mm porcelain veneer where ∆ TEC < 0; the TZP-porcelain-composite is preloaded in compression and therefore reinforced. This reinforcement advantage becomes smaller as the thickness of the porcelain layer increases. For ∆TEC > 0 the effects are reversed.
W35/9
W35/3
σTEC [MPa]
40
W35/4 W35/11
W35/10
20 0image: chapter_3/Veneer_Porcelain/s_tec.wmf design: chapter_3/Veneer_Porcelain/load_bearing.ppt
-20 Porcelain 1 mm Porcelain 3 mm
-40 -2.0
-1.5
-1.0
-0.5
0.0
0.5
1.0
∆ TEC [10-6 K-1] Fig. 124: Residual stress σTEC due to TEC mismatch between the porcelain and the TZP substrate. ∆TEC = TECporcelain - TECzirconia, grey shaded area: compressive preloading
Appendix
264
8.7±1.0 6.2±0.9
9.3±0.9 6.6±0.8
10.0±0.8 6.2±0.8
6.5±0.7 6.7±0.9
6.7±0.6 7.1±1.0
7.3±0.6 9.7±1.4
6.4±0.7 8.6±1.2
8.1±0.9 9.2±1.5
No.
18
17
16
15
14
13
12
11
No.
48
47
46
45
44
43
42
41
10.6±0.8 6.1±0.9
10.8±0.8 5.8±0.9
6.8±0.7 6.7±1.1
6.6±0.6 7.8±1.1
6.6±0.6 9.8±1.4
5.7±0.6 8.2±1.1
5.1±0.6 7.5±1.3
Upper Jaw
Height
Lower Jaw
Anatomical tooth sizes
10.4±1.0 6.1±0.9
VI.8
Width
Height
Width
Tab. 24: Anatomical height and width of teeth in mm on the right jaw side [12]. Teeth of the left side exhibit equivalent size. The width is determined in the mesialdistal axis and the height from the enamel-dentin line.
VI.9
Prototype Machine
VI.9.1
Principle of the Prototype Machine
The mechanical principle of an integrated digitizing and milling machine (10) is shown in fig. 125. On the right side of the machine the digitizing is performed using either tactile (probe 46, stylos 40) or optical technology. On the left milling is conducted with a milling device (52) which consists of a fine milling spindle (54, tool 58) and of a rough milling spindle (52, tool 60). The framework model (44) and the blank (42) are moved by means of one translatory three-axis stage (16) and are rotated by 180° around
265
Appendix
one axis (40) using rotatory device (36). A barcode scanner (51) reads the enlargement factor of the blank to be machined. All functions of the machine are controlled by means of a personal computer (50). The ground plate (12) and the portal frame (14) assure the positions of all devices.
image: figures/machine_prototype/machine_principle.jpg
Fig. 125: Principle of an machine for an automated fabrication of all-ceramic dental frameworks [13].
Appendix
VI.9.2
(a)
266
Prototype Machines
(b)
Image: machine_prototype/prototypes_01.jpg
(c)
Fig. 126: Prototype machines for an integrated digitizing and milling of ceramic frameworks. (a, b) tactile (Renishaw), (c) conoscopical digitizing (Optimet).
267
Appendix
VI.10 Series Machine
Image: figures/Cercon_Geraete.jpg
Fig. 127: Series machine built on the same principles as the prototypes (Courtesy of Degussa Dental, Hanau, Germany)
Appendix
268
VI.11 References [1]
Gregg, S.J. and Sing, K.S.W.: Adsorption, Surface Area, and Porosity., 1982.
[2]
Wachtman, J.B.: "Mechanical Properties of Ceramcis.". 1996, New York, Chichester, Brisbane, Toronto, Singapore: John Wiley & Sons, Inc.
[3]
Evans, A.G. and Charles, E.A.: "Fracture Toughness Determinations by Indentation." Journal of the American Ceramic Society, 1976. 59(7-8): p. 371-372.
[4]
Binner, J.G.P. and Stevens, R.: "The Measurement of Toughness by Indentation."
British Ceramic Transaction Jounal, 1984. 83: p. 168-172. [5]
Ostojic, P. and McPherson, R.: "A Review of Indentation Fracture Theory: its Development, Principles and Limitations." International Journal of Fracture, 1987. 33: p. 297-312.
[6]
Anstis, G.R., Chantikul, P., Lawn, B.R., and Marshall, D.B.: "A Critical Evaluation of Indentation Techniques for Measuring Fracture Toughness: I, Direct Crack Measurements." Journal of the American Ceramic Society, 1981. 64(9): p. 533538.
[7]
Chantikul, P., Anstis, G.R., Lawn, B.R., and Marshall, D.B.: "A Critical Evaluation of Indentation Techniques for Measuring Fracture Toughness: II, Strength Method." Journal of the American Ceramic Society, 1981. 64(9): p. 539-543.
[8]
Niihara, K., Morena, R., and Hasselman, D.P.H.: "Evaluation of KIc of brittle solids by the indentation method with low crack-to-indent ratios." Journal of Materi-
als Science, 1982(1): p. 13-16. [9]
Michel, B.: "Korngrenzenglasphasen in tetragonalem Zirkonoxid.", PhD, 1993, in
Departement of Materials, Institute for Nonmetallic Inorganic Materials. Swiss Federal Institute of Technology: Zurich. [10] Weast, R.C., Lide, D.R., Astle, M.J., and Beyer, W.H. ( eds.): "CRC Handbook of Chemistry and Physics." 70th ed. 1989, CRC Press, Inc.: Boca Raton, Florida. [11] Dorsch, P.: "Messmethode zur Spannungsberechnung in Metall/Porzellan-Verbundkörpern." Berichte der deutschen keramischen Gesellschaft, 1979(11-12): p. 328331.
269
Appendix
[12] Schumacher, G.-H.: "Odontographie: Eine Oberflächenanatomie der Zähne.", 4. ed. 1983, Leipzig: Johann Ambrosius Barth. [13] Filser, F., Kocher, P., Gauckler, L.J., Lüthy, H., and Schärer, P.: "Werkzeugmaschine zur Herstellung von Grundgerüsten für Zahnersatz.", 2001, European Patent Office 2001/24 (eurpäisches Patentblatt, EP 1106146.
Appendix
270
271
VII.
List of Figures
List of Figures
1.
Dental ceramic materials properties used for restorations. . . . . . . . . . . . . . . . . . . . 7
2.
Direct Ceramic Machining (DCM) system approach. . . . . . . . . . . . . . . . . . . . . . 10
3.
Steps for the fabrication of blanks. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 44
4.
Primary particles and spray dried granules of TZ-3YB. . . . . . . . . . . . . . . . . . . . 47
5.
Monomodal agglomerate size distribution for TZ-3YB and TZ-3YB-E powder. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 49
6.
Materials for wet-bag-CIP and fabricated shapes and sizes of green bodies. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 51
7.
Density of TZ-3YB green bodies as function of the compaction pressure. . . . . . 53
8.
Pore radius distribution in the TZ-3YB green body for different compaction pressures. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 54
9.
Pore size distribution and mean pore radius for TZ-3YB green body. . . . . . . . . . 55
10.
Mean pore radius as function of the compaction pressure. . . . . . . . . . . . . . . . . . 55
11.
Vickers hardness as a function of the compaction pressure. . . . . . . . . . . . . . . . . 56
12.
DTA / TG for TZ-3YB and TZ-3YB-E. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 58
13.
Dilatometry of TZ-3YB-E green bodies using different heating rates. . . . . . . . . 59
14.
Thermal behaviour of TZ-3YB-E. Comparison between DTA and dilatometry. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 60
15.
Relative length change of the TZ-3YB-E compacts after heat treatment at different maximum temperatures. . . . . . . . . . . . . . . . . . . . . . . . . . . . 61
16.
Density as a function of compaction pressure and heat treatment temperature for TZ-3YB specimen. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 63
17.
Local density distribution in a blank. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 65
18.
Pore radius distribution in TZ-3YB compacts for different temperatures. . . . . . . 66
List of Figures
272
19.
Derivation of the cumulative volume for blanks fabricated with 300 MPa compaction pressure. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 67
20.
Pore radius distribution in a blank. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 68
21.
Mean pore radius of blanks as function of compaction pressure and heat treatment. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 68
22.
Fracture surface of a blank presintered at 850 °C for 120 min. . . . . . . . . . . . . . 69
23.
Vickers Hardness of blanks as function of the heat treatment temperature and the compaction pressure. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 71
24.
Machinability as a function of the presintering temperature. . . . . . . . . . . . . . . . 74
25.
Machinability as a function of the compaction pressure. . . . . . . . . . . . . . . . . . . 75
26.
Chippings produced by milling a blank using tool of different diameters. . . . . . 77
27.
Chippings produced by milling a green body using a finish milling tool. . . . . . 79
28.
Chippings produced by milling a green body using a rough milling tool. . . . . . 80
29.
Model for the different mechanisms of material removal. . . . . . . . . . . . . . . . . . 81
30.
Reflection spectra for TZP infiltrated with different cations. . . . . . . . . . . . . . . . 84
31.
Reflection spectra for praseodym in TZP as function of its concentration. . . . . 85
32.
Reflection spectra for iron in TZP as function of its concentration. . . . . . . . . . . 86
33.
Reflection spectra for vanadium in TZP as function of its concentration. . . . . . 86
34.
Colour of sintered TZP containing different concentrations of Pr3+, Fe3+, and V4+ cations. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 87
35.
Colour of TZP infiltrated with equivalent cation concentrations as function of sintering temperature. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 88
36.
Shrinkage and shrinkage rate for TZP blanks with different dopants. . . . . . . . . 90
37.
Penetration depth of different dopants for a constant storage time. . . . . . . . . . . 92
38.
Reflection spectra as a function of the thickness of two different porcelains on TZP substrates. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 94
39.
L*a*b* analysis as a function of the coating thickness of two different porcelains. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 95
40.
Analysis of blanks and green bodies fabricated from TZ-3YB and TZ-3YB-E powder. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 98
41.
Variation of density, shrinkage factor, and enlargement factor of fabricated blanks. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 100
273
List of Figures
42.
Grinding of TZP test bars. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 106
43.
Monomodal particle size distribution of W35/11 powder. . . . . . . . . . . . . . . . . . 109
44.
Fabrication of bilayered TZP-veneer-test bars. . . . . . . . . . . . . . . . . . . . . . . . . . 111
45.
Mechanical system for a bilayered bar in three-point bending. . . . . . . . . . . . . . 113
46.
Bend strength of bilayered TZP-porcelain structures as function of TEC difference for two veneer porcelain thicknesses. . . . . . . . . . . . . . . . . . 115
47.
Resulting stress in the porcelain’s edge fibre of the bilayer bar. . . . . . . . . . . . . 116
48.
Failure of a TZP - W35/11 bilayer with the veneer in tensile mode. . . . . . . . . . 117
49.
Bend strength of 4 mm-bilayers consisting of TZP and W35/11 veneer of varying thickness ratios. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 118
50.
Comparison of the measured to the calculated load bearing capacity of bilayers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 120
51.
Weibull plot for test bars of constant overall thickness as function of the TZP ratio and of the test mode. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 121
52.
Steps of the DCM process to fabricate all-ceramic dental restorations. . . . . . . . 125
53.
Scheme of a three-unit bridge framework with definitions of the terms. . . . . . . 127
54.
Plastic frameworks of crowns and bridges as fabricated by the dental technician. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 128
55.
Setup for digitizing. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 130
56.
Clamping device for digitizing a framework model. . . . . . . . . . . . . . . . . . . . . . 131
57.
Digitizing pathway for a three-unit bridge framework. . . . . . . . . . . . . . . . . . . . 132
58.
Different types of styli for tactile digitizing. . . . . . . . . . . . . . . . . . . . . . . . . . . . 133
59.
Digitizing a horizontal to vertical transistion using different digitizing styli. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 134
60.
Plastic deformation of metal reducing the accuracy of digitizing. . . . . . . . . . . . 135
61.
Distortion at abrupt transistions caused by too high digitizing speed. . . . . . . . . 137
62.
Point clouds by tactile and by optical digitizing. . . . . . . . . . . . . . . . . . . . . . . . . 139
63.
System setup for milling. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 147
64.
Clamping device for machining a blank. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 148
65.
Tool travel pathway for machining of a three-unit framework. . . . . . . . . . . . . . 149
66.
Different milling tools for finishing and roughing. . . . . . . . . . . . . . . . . . . . . . . 150
List of Figures
274
67.
Enlarged porous ceramic framework in comparison to the original model. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 152
68.
Rough milling process of a three-unit framework. . . . . . . . . . . . . . . . . . . . . . . 153
69.
Fine milling process of a three-unit framework. . . . . . . . . . . . . . . . . . . . . . . . . 154
70.
Edge quality for conventional and climb milling type. . . . . . . . . . . . . . . . . . . . 155
71.
Cervical edges after milling in presintered state. . . . . . . . . . . . . . . . . . . . . . . . 156
72.
Inspection of the tool tip after being in service. . . . . . . . . . . . . . . . . . . . . . . . . 157
73.
High temperature furnaces used for sintering dental restorations. . . . . . . . . . . 160
74.
Dimensional accuracy and homogeneity of shrinkage determined using a three-unit bridge on its master cast. . . . . . . . . . . . . . . . . . . . . . . . . . . . 162
75.
Homogeneity of shrinkage and dimensional accuracy determined using a calibration specimen. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 163
76.
Shrinkage and sinter rate as function of the temperature. . . . . . . . . . . . . . . . . . 165
77.
Start of sintering and maximum sinter rate as a function of the compaction pressure or green density. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 167
78.
Temperature at max. sinter rate and total linear shrinkage as function of the compaction pressure. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 168
79.
Microstructure of sintered TZ-3YB. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 169
80.
Microstructure of sintered TZ-3YB-E. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 170
81.
Grain size analysis for sintered TZ-3YB and TZ-3YB-E. . . . . . . . . . . . . . . . . 171
82.
Vickers hardness and fracture toughness of TZ-3YB. . . . . . . . . . . . . . . . . . . . 173
83.
Vickers indentation in TZP showing the Palmqvist cracks. . . . . . . . . . . . . . . . 174
84.
Influence of the sintering temperature and the dwell time on the density of TZ-3YB. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 175
85.
Translucency of TZ-3YB as a function of the sample’s thickness. . . . . . . . . . . 177
86.
Transmittance as function of theTZ-3YB sample thickness. . . . . . . . . . . . . . . 178
87.
Dimensional accuracy and shrinkage homogeneity of sintered three-unit framework. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 179
88.
Example of a five-unit framework. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 180
89.
Variation of the density of sintered frameworks. . . . . . . . . . . . . . . . . . . . . . . . 181
90.
Coating the zirconia framework with veneer porcelain. . . . . . . . . . . . . . . . . . 186
275
List of Figures
91.
Clinical fracture of an all-ceramic bridge. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 192
92.
Design of the frameworks for load bearing and reliability investigations. . . . . 194
93.
W35/11 porcelain thickness on TZP framework. . . . . . . . . . . . . . . . . . . . . . . . . 195
94.
Framework test setup in diagram and in reality. . . . . . . . . . . . . . . . . . . . . . . . . 196
95.
FEA for the framework test setup. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 198
96.
Principal stress distribution in the mesial-distal cross-section of a loaded framework. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 199
97.
Principal stress distribution in the loaded framework from gingival side. . . . . . 200
98.
Minimum, maximum and average load for bridge test with TZP, In-Ceram and Empress2 frameworks. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 201
99.
Load bearing capacity and reliability as a function of the frameworks’ material. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 202
100. Load bearing capacity and reliability as a function of the TZ-3YB frameworks’ connector cross-section area. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 203 101. Load bearing capacity and reliability as a function of the veneer thickness at connector’s gingival side. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 204 102. Load bearing capacity and reliability of TZ-3YB and In-Ceram test bridges. . . 205 103. Load-displacement diagram of TZP frameworks. . . . . . . . . . . . . . . . . . . . . . . . 206 104. Schematic diagrams of load-displacement behaviour of TZP and In-Ceram test bridges. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 207 105. Typical fracture of the unveneered and veneered TZP framework. . . . . . . . . . . 208 106. Fracture surfaces of TZP, In-Ceram, and Empress2 frameworks. . . . . . . . . . . . 209 107. Fracture surfaces of TZP and In-Ceram test bridges. . . . . . . . . . . . . . . . . . . . . . 210 108. Weibull parameters of the test sets. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 212 109. Step-by-step clinical insertion procedure. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 218 110. Tooth preparation guidlines for the dentist. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 219 111. Occlusal tooth reduction of 1.9 to 2.0 mm is required. . . . . . . . . . . . . . . . . . . . 219 112. Avoid grinding the framework in the dense state for shape optimization. . . . . . 220 113. Minimum thickness of the framework is 0.7 mm and of the veneer porcelain is 1.2 mm. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 221 114. Study case at a patient. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 224
List of Figures
276
115. Translucency of an all-ceramic three-unit bridge. . . . . . . . . . . . . . . . . . . . . . . . 226 116. Single crowns for a bruxism patient. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 227 117. A four-unit bridge and a three-unit bridge for a bruxism patient. . . . . . . . . . . . 228 118. Initial situation of a bruxism patient versus the final situation 12 months after insertion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 229 119. Large span-width five-unit bridge. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 230 120. Five-unit bridge with four retainers and one pontic unit and one with three retainers and two pontics. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 231 121. Microstructure of TZ-3Y by Michel. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 258 122. Fabrication of plastic framework model for a three-unit bridge. . . . . . . . . . . . 261 123. Stress in the porcelain edge fibre. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 262 124. Residual stress due to TEC mismatch between the porcelain and the TZP substrate. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 263 125. Principle of an machine for an automated fabrication of all-ceramic dental frameworks. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 265 126. Prototype machines for an integrated digitizing and milling of ceramic frameworks. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 266 127. Series machine built on the same principles as the prototypes. . . . . . . . . . . . . 267
277
List of Tables
VIII. List of Tables 1.
Comparison of ceramic forming methods enabling a fabrication of blanks. ........ 43
2.
Average chemical composition of TZ-3YB and its variance. .............................. 46
3.
TZ-3YB and TZ-3YB-E powder and sintered part characteristics. ...................... 48
4.
pH-values of the coloration solutions for different cations and concentrations. ...................................................................................................... 83
5.
Density tolerances in the fabrication process of the blanks. ................................. 99
6.
Shrinkage and enlargement factor tolerances in the fabrication process of the blanks. ......................................................................................................... 99
7.
Chemical composition and TEC of the veneer porcelains. ................................ 108
8.
Firing cycles for TZP test bars coated with veneer porcelain. ........................... 110
9.
Weibull modulus and characteristic bend strength of the bilayered test bars. ... 122
10.
Shrinkage and enlargement factor tolerances in the fabrication process of the blanks. ....................................................................................................... 145
11.
Specification of the milling tools. ....................................................................... 151
12.
Firing cycles for applying the veneer porcelain on TZP frameworks. ............... 185
13.
Overview on the number of fabricated specimen for load bearing and reliability analyses. ............................................................................................. 195
14.
Materials data as used for FEA. .......................................................................... 198
15.
Weibull parameters of the framework specimen. ............................................... 205
16.
Overview of the results of the clinical study after the first recall. ...................... 223
17.
Chemical analysis of used lots of TZ-3YB powder. .......................................... 251
18.
Chemical analysis of used lots of TZ-3YB-E powder. ....................................... 251
19.
Assumptions for the variables for error calculating of the enlargement factor. ............................................................................................. 254
20.
Errors in enlargement factor’s equations. ........................................................... 255
List of Tables
278
21.
Assumptions for the error calculation of the bend strenght. .............................. 256
22.
Assumptions for the variables for error calculation of the density by Archimedes. .................................................................................................. 256
23.
Absolute and relative errors for the fracture toughness. ..................................... 258
24.
Anatomical height and width of teeth. ............................................................... 264
279
IX. CT: CAD: CAM: CBN: CIP: DCC: DCM: DTA: d50: HSS: ISO: NC: MOR: PCD: PFM: PVA: RP: RT: SEM: SFF: TD: TG: TZP: TZ-3Y: TZ-3YB: TZ-3YB-E: Y-TZP: XRD:
Abbreviations
Abbreviations
Computer Tomography Computer Aided Design Computer Aided Manufacturing Cubic Boron Nitride Cold Isostatic Pressing Direct Coagulation Casting Direct Ceramic Machining Differential Thermo Analysis Median Grain Size High-Speed-Steel International Organization for Standardization Numerical Controlled Modulus of Rupture Polycrystalline Diamond Porcelain-Fused-to-Metal Poly-Vinyl-Acryl (Binder) Rapid Prototyping Room Temperature Scanning Electron Microscopy Solid Freeform Fabrication Theoretical Density Thermogravimetry Tetragonal Zirconia Polycrystals Tetragonal Zirconia Polycrystals with 3 mol-% Yttria Tetragonal Zirconia with 3 mol-% Yttria and binder (Tosoh grade) Tetragonal Zirconia with 3 mol-% Yttria, 0.25 wt-% Al2O3 and binder (Tosoh grade) TZP stabilized with yttria X-ray diffraction
Abbreviations
280
281
X.
Curriculum Vitae
Curriculum Vitae Personal Data
Name
Frank Thomas Filser
Address
Winkelrainweg 15, 8102 Oberengstringen, Switzerland
Birth
April 24th, 1966 in St. Wendel, Germany
Citizenship
German
Marital Status
unmarried
Education
2001
PhD at the Swiss Federal Institute of Technology, Nonmetallic Inorganic Materials, Department of Materials, Zürich (Switzerland)
1992
Diploma in Mechanical Engineering at the University of Kaiserslautern (Germany)
1985
German general matriculation at the secondary school in Kusel (Germany)
Work Experience
1996 - present
Swiss Federal Institute of Technology; Institute of Nonmetallic Inorganic Materials, Dep. of Materials, head Prof. Dr. L.J. Gauckler
1992 - 1995
University of Kaiserslautern; chair of Manufacturing and Production Engineering (Lehrstuhl für Fertigungstechnik und Betriebsorganisation) and the CIM-Center Kaiserslautern, head Prof. Dr.-Ing. G. Warnecke
1986 - 1992
Euro Consult
Curriculum Vitae
282
283
XI. XI.1
Work
Work
Publications
❍ F. Filser, P. Kocher, F. Weibel, H. Lüthy, P. Schärer, and L.J. Gauckler: "Reliability and Strength of All-Ceramic Dental Restorations fabricated by Direct Ceramic Machining (DCM). Zuverlässigkeit und Festigkeit vollkeramischen Zahnersatzes im DCM-Verfahren." International Journal of Computerized Dentistry, 2001. 4: p. 83-106. ❍ B. Sturzenegger, A. Fehér, H. Lüthy, M. Schumacher, O. Loeffel, F. Filser, P. Kocher, L.J. Gauckler, P. Schärer: "Klinische Studie von Zirkonoxidbrücken im Seitenzahngebiet hergestellt mit dem DCM-System", Acta Med Dent Helv 5: 131139 (2000). ❍ U.P. Schönholzer, F. Filser, P. Kocher, L.J. Gauckler : “Comparison of Processing Methods for the Fabrication of a Surface Pattern in Zirconia”, The American Ceramic Society Bulletin 12 (2000), p. 45-47 ❍ M. Estermann, H. Reifler, W. Steurer, F. Filser, P. Kocher, L.J. Gauckler: "A hightemperature furnace for X-ray diffraction with directly machined alpha-Al2O3 ceramic part", Journal of Applied Crystallography (1999), 32, p. 833-836. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Direct Ceramic Machining (DCM) for Production of All-Ceramic Dental Bridges”, ECerS 1999, 6th Conference and Exhibition of the European Ceramic Society, 20 - 24th June 1999, Brighton, UK, British Ceramic Proceedings No. 60, Volume 2, p. 527-528. ❍ H. Lüthy, F. Filser, L. Gauckler, P. Schärer: „Reliability of DCM Machined Zirconia Bridges: Comparison for Different Interdental Connector Areas.” Journal of Dental Research, vol. 78, 1999, Special Issue, p. 205, abstract no. 793. ❍ F. Filser, M. Eglin, H. Lüthy, P. Schärer, L. Gauckler: „FEA Study of DCM Machined Zirconia vs. Vita-Celay In-Ceram Bridges”. Journal of Dental Research, vol. 78, 1999, Special Issue, p. 204, abstract no. 792. ❍ F. Filser, H. Lüthy, P. Schärer, L. Gauckler: „ All-Ceramic Restorations by New Direct Ceramic Machining Process (DCM)“. Journal of Dental Research, vol. 77, 1998, p. 762, abstract no. 1046. Proceeding: Dental Materials Group, IADR, Microfilm and Microfiche,1998.
Work
284
❍ H. Lüthy, F. Filser, L. Gauckler, P. Schärer: “High Reliable Zirconia Bridges by Direct Ceramic Machining (DCM)”. Journal of Dental Research, vol. 77, 1998, p. 762, abstract no. 1045. Proceeding: Dental Materials Group, IADR, Microfilm and Microfiche,1998. ❍ F. Filser, H. Lüthy, P. Schärer, L. Gauckler: „ All-Ceramic Dental Bridges by Direct Ceramic Machining (DCM)“. In: Materials in Medicine, Materials Day, Department of Materials, Ed. M.O. Speidel, P.J. Uggowitzer, vdf Hochschulverlag AG, ETH Zürich, Zürich, May 1998, p. 165-189. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer and L.J. Gauckler: “All-Ceramic Dental Bridges by the Direct Ceramic Machining Process (DCM)“. Bioceramics, Volume 10, Ed. L. Sedel, C. Rey, Elsevier Science Ltd, Proceedings of the 10th International Symposium on Ceramics in Medicine, Paris, France, 5-9 October 1997, p 433-436. ❍ Y. Kakehashi, H. Lüthy, F. Filser, L. Gauckler and P. Schärer: “Strength of Zirconia/ Veneer Bilayered Structures”. Journal of Dental Research, vol. 76, 1997, p. 138, abstract no. 996. Proceeding: Dental Materials Group, IADR, Microfilm and Microfiche,1997. ❍ G. Warnecke, M. Radtke, F. Filser: “Produktmodelle als Grundlage vernetzter Produktentwicklungsprozesse. Unternehmens- und prozeßspezifische Produktmodelle.” In: wt Produktion und Management 85 (1995) 3, S. 132-136 ❍ F. Filser, M. Radtke, C. Schulz: “Ein Beitrag zur Modellbildung in der rechnergestützten Arbeitsplanung”. LSA 95-03, Zentrum für lernende Systeme und Anwendungen (LSA), Universität Kaiserslautern, 1995 ❍ G. Warnecke, C. Schulz, F. Filser: “A Reference Model for the Development of Computer-Aided Process Planning Systems”. In: Production Engineering I/2 (1994), S. 169-172 ❍ M.M. Richter, K.-D. Althoff, F. Filser, A. de la Ossa, J. Paulokat, R. Präger, S. Wess: “Teilprojekt X 9: Lernen und Analogie in technischen Expertensystemen”. In: P. Deussen (Hrsg.): SFB 314 - Arbeits- und Ergebnisbericht für die Jahre 1991 1992 - 3/1993, Universität Karlsruhe, 1993
XI.2
Thesis during Study
❍ F. Filser: “Konzeptentwicklung eines Ressourcenmodells zur wissensbasierten Arbeitsplanung”. Diplomarbeit, Lehrstuhl Fertigungstechnik und Betriebsorganisation FBK, Universität Kaiserslautern, April 1992
285
Work
❍ F. Filser: “Technischer Modellierer zur Integration des Produktentstehungsprozesses”. Studienarbeit, Lehrstuhl Fertigungstechnik und Betriebsorganisation FBK, Universität Kaiserslautern, Jan. 1991 ❍ F. Filser: “Messung von Dampf-Flüssigkeits-Gleichgewichten formaldehydhaltiger Systeme bei Temperaturen unter 50 °C”. Studienarbeit, Lehrstuhl für Technische Thermodynamik, Universität Kaiserslautern, June 1989
XI.3
Patents
❍ F. Filser, L. Gauckler, P. Kocher, H. Lüthy, P. Schärer: “Zahnkronen und/oder Zahnbrücken”. PCT/CH99/00115, March 16th, 1999. ❍ F. Filser, L. Gauckler, P. Kocher, H. Lüthy, P. Schärer: “Zahnkronen und/oder Zahnbrücken”. EP 943296 A1, Filing: Nov. 11th, 1998, Publication: Sept. 22nd, 1999. ❍ F. Filser, L. Gauckler, P. Kocher, H. Lüthy, P. Schärer: “Verfahren zur Herstellung von Zahnkronen und/oder Zahnbrücken”. EP 943295 A1, Filing: March 17th, 1998, Publication: Sept. 22nd, 1999. ❍ F. Filser, P. Kocher, L.J. Gauckler, H. Lüthy, P. Schärer, and S. Fecher, H. Hörhold, P. Kreuder (Degussa Dental): “Haltevorrichtung für ein Zahnersatz- oder Grundgerüstmodell”. CH 2000 2415 /00, Filing: Dec. 12th, 2000 ❍ F. Filser, P. Kocher, L.J. Gauckler, H. Lüthy, P. Schärer, and S. Fecher, H. Hörhold, P. Kreuder (Degussa Dental): “Haltevorrichtung für einen Keramikrohling”. CH 2000 2388/00, Filing: Dec. 7th, 2000
XI.4
Presentations
❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Direct Ceramic Machining (DCM) for Production of All-Ceramic Dental Bridges”, ECerS 1999, 6th Conference and Exhibition of the European Ceramic Society, June 24th, 1999, Brighton, UK. ❍ L.J. Gauckler, F. Filser, P. Kocher, H. Lüthy, P. Schärer, A. Fehér: "High Strength, High Reliable All-Ceramic Dental Bridges", International Symposium on "Advanced Materials with Biomedical Applications", NIST (National Institute of Standards and Technology), June 7-8, 1999, Gaithersburg, Maryland, USA. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Direct Ceramic Machining (DCM) for Production of All-Ceramic Dental Bridges”, AmCerS 1999, 101st
Work
286
Annual Meeting of the American Ceramic Society, April 25 - 28th, 1999, Indianapolis (IN), USA. ❍ F. Filser: “All-Ceramic Dental Bridges by the new Direct Ceramic Machining Process”, D-Werk Kolloquium, Zürich, April 14th, 1999. ❍ H. Lüthy, F. Filser, L. Gauckler, P. Schärer: „Reliability of DCM Machined Zirconia Bridges: Comparison for Different Interdental Connector Areas.” 77th General Session & Exhibition of the IADR, Vancouver, Canada, March 10-13th, 1999. ❍ F. Filser, M. Eglin, H. Lüthy, P. Schärer, L. Gauckler: „FEA Study of DCM Machined Zirconia vs. Vita-Celay In-Ceram Bridges”. 77th General Session & Exhibition of the IADR, Vancouver, Canada, March 10-13th, 1999. ❍ F. Filser, H. Lüthy, P. Schärer, L. Gauckler: „All-Ceramic Restorations by New Direct Ceramic Machining Process (DCM)“. 76th General Session and Exhibition of the International Association for Dental Research, June 24-27, 1998, Nice, France. ❍ H. Lüthy, F. Filser, L. Gauckler, P. Schärer: “High Reliable Zirconia Bridges by Direct Ceramic Machining (DCM)”. 76th General Session and Exhibition of the International Association for Dental Research, June 24-27, 1998, Nice, France. ❍ F. Filser, H. Lüthy, P. Schärer, L.J. Gauckler: “All-Ceramic Dental Bridges and Implants”. Materials Day, Zürich, May 29th, 1998. ❍ F. Filser, H. Lüthy and L. Gauckler: „All-Ceramic Dental Bridges“. PPM-Module Meeting (Swiss Priority Program on Materials Research), presentation of the project 4.2C, Lausanne, November 6th, 1997. ❍ F. Filser , P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “All-Ceramic Dental Bridges by the Direct Ceramic Machining Process (DCM)“. 10th International Symposium on Ceramics in Medicine, Paris, France, 8. October 1997. ❍ F. Filser, H. Lüthy and L. Gauckler: “All-Ceramic Dental Bridge Restorations“. PPM-Module 4 Meeting (Swiss Priority Program on Materials Research), Bern, November 7th, 1996. ❍ Y. Kakehashi, H. Lüthy, F. Filser, L.Gauckler and P.Schärer: “Strength of zirconia/ veneer bilayered structures”. 75th General Session and Exhibition of the International Association for Dental Research, 19.-23. March 1997, Orlando FL, USA.
XI.5
Posters
❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Materials Selection for DCM All-Ceramic Dental Bridges”. Ausstellung von Firmen und Forschungsinstitutionen, Materialien für eine bessere Zukunft, Tag der offenen Tür, Departement Werkstoffe, ETH Zürich, Sept. 24th -25th, 1999.
287
Work
❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “All-Ceramic Dental Bridges by Direct Ceramic Machining (DCM)”. Ausstellung von Firmen und Forschungsinstitutionen, Materialien für eine bessere Zukunft, Tag der offenen Tür, Departement Werkstoffe, ETH Zürich, Sept. 24th -25th, 1999. ❍ F. Filser, A. Fehér, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Clinical Application of DCM Machined All-Ceramic Dental Bridges”. Ausstellung von Firmen und Forschungsinstitutionen, Materialien für eine bessere Zukunft, Tag der offenen Tür, Departement Werkstoffe, ETH Zürich, Sept. 24th -25th, 1999. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Materials Selection for DCM All-Ceramic Dental Bridges”. AmCerS 1999, 101st Annual Meeting of the American Ceramic Society, April 25 - 28th, 1999, Indianapolis (IN), USA. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “All-Ceramic Dental Bridges by Direct Ceramic Machining (DCM)”. AmCerS 1999, 101st Annual Meeting of the American Ceramic Society, April 25 - 28th, 1999, Indianapolis (IN), USA. ❍ F. Filser, A. Fehér, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Clinical Application of DCM Machined All-Ceramic Dental Bridges”. AmCerS 1999, 101st Annual Meeting of the American Ceramic Society, April 25 - 28th, 1999, Indianapolis (IN), USA. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Materials Selection for DCM All-Ceramic Dental Bridges”. Swiss Physical Society Annual Meeting 1999, Feb. 26th, 1999, Bern. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “All-Ceramic Dental Bridges by Direct Ceramic Machining (DCM)”. Swiss Physical Society Annual Meeting 1999, Feb. 26th, 1999, Bern. ❍ A. Fehér, F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Clinical Application of DCM Machined All-Ceramic Dental Bridges”. Swiss Physical Society Annual Meeting 1999, Feb. 26th, 1999, Bern. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “All-Ceamic Dental Bridges by Direct Ceramic Machining”, PPM Fall Meeting 1998 (Swiss Priority Program on Materials Research), Nov. 13th, 1998, Bern. ❍ A. Fehér, F. Filser, H. Lüthy, L.J. Gauckler, P. Schärer: “Clinical Application of DCM Machined Bridges”. PPM Fall Meeting 1998 (Swiss Priority Program on Materials Research), Nov. 13th, 1998, Bern. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “Materials Selection for All-Ceamic Dental Bridges”. PPM Fall Meeting 1998 (Swiss Priority Program on Materials Research), Nov. 13th, 1998, Bern. ❍ F. Filser, P. Kocher, H. Lüthy, P. Schärer, L.J. Gauckler: “All-Ceamic Dental Bridges by Direct Ceramic Machining”. Materials Day: Materials in Medicine, ETH Zürich, May 29th, 1998, Zürich.
Work
288
❍ A. Fehér, F. Filser, H. Lüthy, L.J. Gauckler, P. Schärer: “Clinical Application of DCM Machined Bridges”. Materials Day: Materials in Medicine, ETH Zürich, May 29th, 1998, Zürich. ❍ F. Filser, H. Lüthy, P. Kocher, O. Loeffel, M. Schumacher and L.J. Gauckler: “All Ceramic Dental Bridges by Direct Ceramic Machining (DCM)”. Biosurf (Biomaterials: Surfaces and Biocompatibility), Sept. 25 -26th, 1997, Zürich.
XI.6
Moderation
❍ “Knowhow Transfer in Switzerland”. Round table at the PPM-Meeting, Materials Innovation and Knowledge Transfer for Swiss Industry, Interlaken, Nov. 11th and 12th, 1999.
