deyar jallal hadi mahmood - muep mahmood2... · deyar jallal hadi mahmood on core and bi-layered...

DE

YA

R M

AH

MO

OD

M

ALM

Ö U

NIV

ER

SIT

Y 2

01

5O

N C

OR

E A

ND

BI-LA

YE

RE

D A

LL-CE

RA

MIC

FIXE

D D

EN

TAL P

RO

STH

ES

ES

, DE

SIG

N A

ND

ME

CH

AN

ICA

L PR

OP

ER

TIES

DO

CT

OR

AL

DIS

SE

RTA

TIO

N IN

OD

ON

TO

LOG

Y

DEYAR JALLAL HADI MAHMOODON CORE AND BI-LAYERED ALL -CERAMIC FIXED DENTAL PROSTHESES, DESIGN AND MECHANICAL PROPERTIESStudies on stabilized zirconiumdioxide

O N C O R E A N D B I - L A Y E R E D A L L - C E R A M I C F I X E D D E N T A L P R O S T H E S E S , D E S I G N A N D M E C H A N I C A L P R O P E R T I E S

Malmö University Faculty of Odontology Doctoral Dissertations 2015

© Deyar Jallal Hadi Mahmood, 2015

Photographs and illustrations: Deyar Jallal Hadi Mahmood

ISBN 978-91-7104-405-1 (print)

ISBN 978-91-7104-406-8 (pdf)

Holmbergs, Malmö 2015

DEYAR JALLAL HADI MAHMOOD ON CORE AND BI-LAYERED ALL -CERAMIC FIXED DENTAL PROSTHESES, DESIGN AND MECHANICAL PROPERTIES

Malmö University, 2015Department of Prosthetic Dentistry

Faculty of Odontology Malmö, Sweden

Studies on stabilized zirconiumdioxide

This publication is also available in electronic format at: http://dspace.mah.se/handle/2043/18473

To my entire family.

Acquire knowledge, and learn tranquility and dignity

Umar ibn al-Khattab

TABLE OF CONTENTS

LIST OF PUBLICATIONS ................................................ 11

THESIS AT A GLANCE ................................................... 12

ABSTRACT ................................................................. 13

POPULÄRVETENSKAPLIG SAMMANFATTNING ................. 15

ABBREVIATIONS AND DEFINITIONS .............................. 18

INTRODUCTION ......................................................... 19Treatment planning (Clinical application of dental ceramics) ......19Dental Ceramics ..................................................................20

Ceramic classification ......................................................20Dental glass-based/porcelain systems ...............................21Glass ceramics ................................................................21Glass-infiltrated oxide-ceramics (hybride-ceramics) ...............23Oxide ceramics ...............................................................23Stabilized Zirconium Dioxide ...........................................23

Manufacturing Procedure ......................................................25CAD/CAM .....................................................................26

Effects of the design of all-ceramic dental restorations and in-vitro test setup ...........................................................27

FDP core design ..............................................................27In-vitro test setup for all-ceramic FDPs ..................................28

Follow up and maintenance of dental ceramic restorations ......29What do we know today? Final remarks ................................30

AIMS ......................................................................... 31Specific aims ......................................................................31

MATERIALS AND METHODS ........................................... 33Laboratory Procedures ..........................................................33 Study I ............................................................................33 Study II ...........................................................................38 Study III ...........................................................................40 Study IV .........................................................................42

RESULTS .................................................................... 51Study I ...............................................................................51Study II ...............................................................................53Study III ...............................................................................55Study IV ..............................................................................56

Statistical Analyses ...........................................................61

DISCUSSION .............................................................. 62Methods: In-vitro studies and test setup ...................................62

Choice of specimen design ...............................................62CAD/CAM system ...........................................................63Veneer Techniques ..........................................................65Supporting tooth analogues ..............................................65Cementation and water storage .........................................66Artificial aging (heat treatment, thermocycling, mechanical preload) ........................................................67Load-to-fracture ...............................................................68

Discussion of results .............................................................69Intraoral loads and mechanical preload..............................69

FDP design .........................................................................70The connector/radius ......................................................70Milling dental ceramics ....................................................74LTD ................................................................................75

Veneering techniques and fractures ........................................76Veneer failure predictions .....................................................79

Multifaceted factors .........................................................79Thermal stresses .............................................................79Inappropriate veneering thickness ....................................80Defects in the veneering material ......................................81Multifaceted bond failure ..................................................81Choice of tooth analogues ...............................................81

Fracture mode analyse .........................................................82Clinical significance .............................................................83Future investigations .............................................................83

CONCLUSIONS .......................................................... 85

ACKNOWLEDGEMENTS ............................................... 87

REFERENCES ............................................................... 89

PAPERS I – IV ..............................................................103

11

LIST OF PUBLICATIONS

This thesis is based on the following articles, which will be referred to in the text by their Roman numerals. The articles are appended at the end of the thesis.

I. Bahat Z, Mahmood DJ, Vult von Steyern P. Fracture strength of three-unit fixed partial denture cores (Y-TZP) with different connector dimension and design. Swed Dent J 2009;33:149-159.

II. Mahmood DJ, Linderoth EH, Vult Von Steyern P. The influence of support properties and complexity on fracture strength and fracture mode of all-ceramic fixed dental prostheses. Acta Odontol Scand. 2011 Jul;69(4):229-37.

III. Mahmood DJ, Linderoth EH, Vult von Steyern P, Wennerberg A. Fracture strength of all-ceramic (Y-TZP) three- and four-unit fixed dental prostheses with different connector design and production history. Swed Dent J 2013;37, 4:179-187.

IV. Mahmood DJ, Linderoth EH, Wennerberg A, Vult von Steyern P. Influence of core design, production technique and material selection on fracture behavior of Y-TZP FDPs produced using different multi-layer techniques: split-file, over-pressing and manually built-up veneers. Submitted.

12

THES

IS A

T A

GLA

NC

E

Stud

yA

imM

etho

dsIll

ustr

atio

n M

ain

Find

ings

I: Frac

ture

stre

ngth

of t

hree

-un

it fix

ed p

artia

l den

tal

core

s (Y-

TZP)

with

diff

eren

t co

nnec

tor d

imen

sion

and

de

sign

.

Eval

uate

the

effe

ct o

f var

ious

co

nnec

tor d

imen

sion

s an

d w

heth

er

diffe

rent

radi

i of c

urva

ture

in th

e em

bras

ure

area

of t

he c

onne

ctor

co

uld

affe

ct th

e fra

ctur

e str

engt

h of

th

ree-

unit

all-c

eram

ic fi

xed

dent

al

pros

thes

es (F

DPs

) mad

e of

Y-T

ZP.

48 th

ree-

unit

FDPs

with

Y-T

ZP

fram

ewor

ks w

ith th

ree c

onne

ctor

di

men

sion

s va

ryin

g fro

m 2

.0

mm

to 3

.0 m

m a

nd tw

o ra

dii o

f gi

ngiv

al em

bras

ure,

0.6

mm

and

0.

9 m

m.

CA

D d

esig

n th

ere-

unit

FDP

Frac

ture

stre

ngth

was

sig

nific

antly

hig

her

with

eac

h in

crea

se in

con

nect

or d

imen

sion

. In

crea

sing

the

radi

us o

f the

gin

giva

l em

bras

ure

from

0.6

mm

to 0

.9 m

m in

the

3 m

m ×

3 m

m c

onne

ctor

gro

up in

crea

sed

fract

ure

stren

gth

by 2

0%.

II:

Influ

ence

of s

uppo

rt pr

oper

ties a

nd co

mpl

exity

on

frac

ture

stre

ngth

and

fra

ctur

e mod

e of a

ll-cer

amic

fix

ed d

enta

l pro

sthes

es.

Eval

uate

the

effe

ct o

f var

ious

toot

h an

alog

ues

used

for i

n-vi

tro s

tudi

es

and

how

thes

e in

fluen

ce te

st re

sults

re

latin

g to

the

fract

ure

stren

gth

of

FDPs

mad

e of

Y-T

ZP.

24 th

ree-

unit

FDPs

with

Y-T

ZP

fram

ewor

ks, te

sted

on th

ree t

ypes

of

toot

h-su

ppor

ting

anal

ogue

m

ade f

rom

alum

inum

, pol

ymer

, or

Dur

aLay

.A

lum

inum

with

Y-T

ZP fr

actu

re

Ther

e sh

ould

be

a sta

ndar

dize

d, s

impl

e te

st se

t-up

whe

n in

-vitr

o te

sting

all-

cera

mic

FD

Ps. T

ooth

ana

logu

es w

ith h

igh

E-m

odul

us

gave

hig

h an

d un

real

istic

load

-at-f

ract

ure

valu

es to

geth

er w

ith a

dver

se fr

actu

re m

odes

co

mpa

red

to F

DPs

invo

lved

in c

linic

al fa

ilure

.

III:

Frac

ture

stre

ngth

of

all-c

eram

ic (Y

-TZP

) thr

ee-

and

four

-uni

t fix

ed d

enta

l pr

osth

eses

with

diff

eren

t co

nnec

tor d

esig

n an

d pr

oduc

tion

histo

ry.

Eval

uate

diff

eren

t CA

D/C

AM

sy

stem

s an

d co

mpa

re c

ore

desi

gns,

incr

easi

ng th

e le

ngth

of

the

FDPs

(i.e

. the

num

ber o

f po

ntic

s) to

det

erm

ine

how

this

af

fect

s th

e fra

ctur

e str

engt

h of

al

l-cer

amic

FD

Ps m

ade

from

Y-T

ZP.

16 th

ree-

unit

and

16 fo

ur-u

nit

FDPs

with

Y-T

ZP fr

amew

orks

an

d va

ryin

g co

nnec

tor d

esig

ns

gene

rate

d by

a m

echa

nica

l sc

anne

r or a

lase

r disp

lace

men

t ga

uge

scan

ner.

Four

-uni

t Y-T

ZP

Def

ault

setti

ngs

of th

e va

rious

CA

D/C

AM

sy

stem

s ha

d a

grea

t im

pact

on

fract

ure

stren

gth.

A c

ruci

al fa

ctor

for l

oad-

bear

ing

capa

city

is th

e de

sign

of t

he ra

dius

of t

he

ging

ival

em

bras

ures

. Inc

reas

ing

the

num

ber

of p

ontic

s fro

m th

ree

to fo

ur d

ecre

ases

the

load

-bea

ring

capa

city

by

near

ly h

alf.

IV:

The

influ

ence

of c

ore

desi

gn, p

rodu

ctio

n te

chni

que

and

mat

eria

l se

lect

ion

on fr

actu

re

beha

vior

of Y

-TZP

FD

Ps

prod

uced

with

diff

eren

t m

ulti-

laye

r tec

hniq

ues;

sp

lit-fi

le, o

ver-p

ress

ing

and

man

ually

bui

lt-up

vene

ers.

Eval

uate

the

effe

ct o

f var

ious

ve

neer

ing

mat

eria

ls, te

chni

ques

, an

d co

re d

esig

ns w

ith r

espe

ct to

th

e fra

ctur

e str

engt

h an

d fra

ctur

e m

ode

of c

urre

ntly

use

d m

ultil

ayer

al

l-cer

amic

Y-T

ZP.

110

thre

e-un

it Y-

TZP

FDPs

with

two f

ram

ewor

k des

igns

: 40

with

laye

red

vene

erin

g 30

with

mill

ed v

enee

ring

10 w

ith p

ress

ed v

enee

ring

30 w

ith c

ore

desig

ns g

ener

ated

by

two

diffe

rent

CA

D/C

AM

sy

stem

s.

Frac

ture

pat

tern

The

desi

gn o

f a fr

amew

ork

is a

cru

cial

fact

or

for t

he lo

ad b

earin

g ca

paci

ty. T

he s

tate

-of-

the-

art d

esig

n ar

e pr

efer

able

sin

ce th

e sp

lit-

file

desi

gned

cor

es c

all f

or a

cro

ss-se

ctio

nal

conn

ecto

r are

a, a

t lea

st 42

% la

rger

, to

have

th

e sa

me

load

bea

ring

capa

city

. Ana

lysi

s of

the

fract

ure

patte

rn s

how

s di

ffere

nces

be

twee

n th

e m

illed

ven

eers

and

ove

r-pre

ssed

or

bui

lt-up

ven

eers

, whe

re th

e m

illed

one

s sh

ow n

umer

ical

ly m

ore

vene

er c

rack

s an

d th

e ot

her g

roup

s on

ly s

how

com

plet

e co

nnec

tor f

ract

ures

.

13

ABSTRACT

Loss of teeth can affect a person’s self-esteem, social life, appearance and oral function. Reconstruction of a missing tooth has scientifically been shown to increase self-esteem and quality of life and to maintain oral function. For many patients a fixed dental prostheses (FDP) is preferred, either tooth- or implant-supported. Improvement and development of all-ceramic materials have made them preferable to other alternatives. However, despite properties of dental ceramics’ well known biocompatibility, good chemical and mechanical, the materials have their weaknesses, such as brittleness and some difficulties with the layering porcelain. Many all-ceramic materials cannot withstand minor flexure; more than 0.1 - 0.3 %, will lead to fracture. Oxide-ceramic, specifically yttria stabilized tetragonal zirconia polycrystals (Y-TZP) has become the most commonly used all-ceramic material. This material has the potential to be used for larger restorations. In addition, one of many challenges is to ensure durable zirconia-based restorations in the oral cavity.

In the clinical situation, crowns and bridges are supported by a combination of different structures with differing properties, i.e. bone, dentine and enamel. The complexity of the supporting tissues in the oral cavity creates stress patterns in the prosthetic material, which need to be considered when designing a dental restoration.

The durability of all-ceramic FDPs is dependent on knowledge of the material and design of the FDPs. In particular the design, shape of the connector and the radius of curvature at the gingival embrasure play a significant role in the load-bearing capacity of FDPs.

The overall aim of this thesis is to evaluate design of zirconia-based restorations in relation to achieving increased fracture resistance. Another aim is related to how the choice of material used for supporting tooth analogues in the test set-up and how this influences test results relating to fracture strength of all-ceramic FDPs.

Study I evaluates different radii (0.60 and 0.90 mm) of curvature in the embrasure of the connector area and different connector dimensions (2 x 2, 3 x 2 and 3 x 3 mm) and their effects on the

14

fracture resistance of 3-unit all-ceramic FPDs made of Y-TZP. The results show that by increasing the radius of the gingival embrasure from 0.6 to 0.9 mm, the fracture strength for a Y-TZP FPD with connector dimension 3 x 3 mm will increase by 20%.

Study II investigated how the choice of material (aluminium, polymer and DuraLay) used for supporting tooth analogues and support complexity influence test results concerning the fracture strength of FDPs made of a brittle material Y-TZP. The outcome of the study demonstrated that Y-TZP FDPs cemented on tooth analogues made of aluminium, with high E-modulus showed a significantly higher load at fracture and a different fracture mode than shown in clinical situations.

Study III evaluates how factors as different default settings for connector design of two different CAD/CAM systems and different radii of curvature in the embrasure area of the connector will affect the fracture strength and the fracture mode of 3-unit, i.e. 4-unit all-ceramic FDPs made from Y-TZP and further to investigate how the number of pontics affect the fracture strength of Y-TZP. The results showed that the most crucial factor for the load-bearing capacity is the design of the radius of the gingival embrasures. Increasing the number of pontics from three to four decreases the load-bearing capacity nearly twice.

Study IV investigate and compare the fracture strength and fracture mode in 11 groups of the currently most used multilayer all-ceramic systems for Y-TZP FDPs, with respect to the choice of core material, veneering material area, manufacturing technique (split-file, over-press, built-up porcelains and glass-ceramics), design of connectors and radius of curvature of FDP cores. The results show that the design of a framework is a crucial factor for the load bearing capacity of an all-ceramic FDP. The state-of-the-art designs are preferable, since the split-file designed cores call for a cross-sectional connector area, at least 42% larger, to have the same load bearing capacity as the state-of-the-art designed cores. Analyses of the fracture patterns demonstrated differences between the milled veneers and over-pressed or built-up veneers, where the milled ones showed numerically more veneer cracks whereas the other groups only showed complete connector fractures. All veneering materials/techniques tested were found, with great safety margin to be sufficient for clinical use both anteriorly and posteriorly.

15

POPULÄRVETENSKAPLIG SAMMANFATTNING

Förlust av tänder kan påverka en persons självkänsla, sociala liv, utseende och orala funktion. Rekonstruktion av en saknad tand har vetenskapligt visat en ökning av självkänsla, livskvalitet och upprätthållande av orala funktioner. För många patienter är en fastsittande konstruktion förstahandsvalet; alternativen är antingen en tandburen eller en implantatstödd tandprotes. Utveckling av helkeramiska material har medfört att protetiska konstruktioner av helkeramisk typ har allt mer blivit tandvårdsteamets förstahandsval. Dentala keramer har, utöver de estetiska fördelarna, flera kemiska och mekaniska egenskaper som talar till deras fördel, men materialen är spröda. Många helkeramiska material är mycket känsliga för böjning; en påkänning större än 0.1 - 0.3%, leder till fraktur. Oxidkeramer, såsom yttriumdioxid-stabiliserad tetragonal polykristallin zirkoniumdioxid (Y-TZP) har ökat i användning och är det vanligaste helkeramiska materialet. Detta material har potential att användas för större tandersättningskonstruktioner. En av många utmaningar är att framställa hållbara, zirconiabaserade tandersättningskonstruktioner. I den kliniska situationen stöds tandersättningskonstruktioner, såsom kronor och broar, av en kombination av olika vävnadstrukturer; käkben, dentin och emalj, med olika mekaniska egenskaper. Komplexiteten hos de stödjande vävnaderna i munhålan skapar spänningsmönster i de dentala materialen, vilket måste beaktas när man designar dentala rekonstruktioner. Hållfastheten hos helkeramiska rekonstruktioner

16

avgörs av dess utformning. Vid design av en brokonstruktion bör särskild vikt läggas vid utformningen av de sammanfogandes balkarna (connectorområdena) och dess krökningsradie vid balkövergångarna, vilken har en avgörande roll för brokonstruktionens bärande förmåga.

Det övergripande målet med föreliggande avhandlingsarbete var att utvärdera hur zirkoniumdioxidbaserade brokonstruktioner bör designas för ökad hållfasthet. Därutöver var syftet att undersöka hur valet av tandanalogsmaterial som används vid laboratorietest och deras utformning, påverkar resultaten avseende hållfasthetvärden och frakturmönster vid mekaniska test av keramiska brokonstruktioner.

Delarbete I utvärderade hur brotthållfastheten hos helkeramiska 3-leds broar, utförda i zirkoniumdioxid påverkas av förändringar i connectorns gingivala radie (0,60 respektive 0,90 mm) samt connectorns dimension (2 x 2, 3 x 2 och 3 x 3 mm). Resultaten visar att genom att öka radien på den gingivala övergången från 0,6 till 0,9 mm, ökar brotthållfastheten med 20 % för 3-leds broar i Y-TZP där connectordimenisonen var 3 x 3 mm.

Delarbete II utvärderade hur valet av material för, och utform-ningen, av tandanaloger avsedda för laboratorietest påverkar resul-tatet avseende hållfasthetvärden och frakturmönster vid mekaniska test av brokonstruktioner i Y-TZP. Resultatet av studien visar att brokonstruktioner cementerade på tandanaloger av material med hög E-modul visade en signifikant högre hållfasthetsvärde, samt att frakturmönstret är av en annan typ än det som erhålls vid kliniska frakturer.

Delarbete III utvärderade hur grundinställningarna avseende connectordesign i två olika CAD/CAM-system påverkar hållfastheten, samt hur en ökad längd på brobalken (dvs. antalet pontics) påverkar brotthållfastheten hos helkeramiska brokonstruktioner tillverkade av Y-TZP. Resultaten visar att den mest avgörande faktorn för brohållfasthet är utformningen av, och radien på, connectorns gingival övergång. En ökning av antalet led (pontics) från tre till fyra minskar brotthållfasthetens med nära hälften.

Delarbete IV utvärderade och jämförde brotthållfastheten hos 3-leds broar med Y-TZP som kärnmaterial. 11 grupper med design och ytkeramer av de vanligast förekommande typerna testades avseende betydelsen av faktorer som connectordesign och gingival radie hos

17

broskelettet samt val av ytkeram och framställningsteknik (fräst, över-pressad, handupplagt porslin). Resultat visade att designen av en brokonstruktion är en avgörande faktor för brotthållfastheten hos en helkeramisk bro. State-of-the-art design är att föredra framför split-file design, som behöver en 42% större connectorarea för att likvärdig brotthållfasthet som state-of-the-art designen uppvisar ska erhållas. Analysen av frakturmönster visar skillnader mellan fräst ytkeram och pressad eller handupplagd ytkeram. Grupperna med frästa ytkeramer uppvisade frakturer i ytkeramen, medan övriga grupper bara uppvisade totalfrakturer. Alla ytkeramiska material/metoder som testades visade stor säkerhetsmarginal som är tillräcklig för den kliniska användet i den orala miljön.

18

ABBREVIATIONS AND DEFINITIONS

Alumina Aluminum oxide

CAD Computer-aided design

CAM Computer-aided manufacturing

CIP Cold isostatic pressing

E-modulus Elastic modulus

FDP Fixed dental prostheses

HIP Hot isostatic pressing

LTD Low temperature degradation

MDP 10-Methacryloyloxydecyl dihydrogen phosphate

MPa Mega Pascal

N Newton

POM-C Polyoxymethylene-copolymer

(m) Stable monoclinic-phase

TC Thermocycling

(t) Tetragonal-phase

Y-TZP Yttria stabilized tetragonal zirconia polycrystals

Yttria Yttrium oxide

Zirconia Zirconium dioxide

19

INTRODUCTION

Loss of teeth can affect a person’s self-esteem, social life, appearance and oral function. According to the WHO’s criteria (2001), tooth loss is a physical impairment because it constitutes loss of important parts of the body (1). Reconstruction of a missing tooth has been scientifically shown to increase self-esteem and quality of life and to maintain oral function (2). The dental team has many options concerning restoration production technique and materials in order to fulfil the patient’s need for aesthetic, functional, long-lasting, and biocompatible restorations. Many patients prefer fixed dental prostheses (FDPs), either tooth- or implant-supported (3).

Treatment planning (Clinical application of dental ceramics)In recent years all-ceramic restorations have increased in popularity because ceramic is an aesthetic and biocompatible material. There is a growing tendency toward replacing metal-based restorations in the posterior and anterior region with all-ceramic materials. Development and improvement of high-strength ceramics and new manufacturing processes have led dental teams to prefer all-ceramic materials for FDP treatment. The success of restorations not only depends on the material selected but also involves personal preferences and prevailing clinical conditions such as tooth preparation, caries, endodontic problems, and the choice of cementation technique (4-6). To avoid stress concentrations, the tooth preparation and the restoration should have smooth, rounded contours. However, the cement-space required between the tooth substance and the FDP material can affect the functionality and quality of the remaining tooth substance.

20

The durability of all-ceramic FDPs depends on clinicians and dental technicians knowledge of the material and design of the FDPs (7). In particular, the design, the shape of the connector, and the radius of the curvature at the gingival embrasure play a significant role in the load-bearing capacity of FDPs (8-11). In clinical situations, crowns and bridges receive support from a combination of different structures with differing E-modulus properties such as bone, dentin, and enamel. The complexity of the supporting tissues in the oral cavity leads to stress patterns that the design of the test setup for an in-vitro study should take into consideration (11, 12).

Dental Ceramics All-ceramic materials have strong covalent and ionic interatomic bonds. They are well-known for their high E-modulus, wear resistance, hardness, and melting temperature, as well as their low thermal expansion coefficient and electrical conductivity. Furthermore, some studies have demonstrated that all-ceramic materials accumulate less plaque than acrylic materials (13-15). All-ceramic materials are more tooth-like than metal-based restorations and have optical properties similar to that of enamel and dentin. Despite dental ceramic’s well-known biocompatibility and good chemical and mechanical properties, the materials have their weaknesses, such as brittleness. Many all-ceramic materials cannot withstand flexure; more than 0.1–0.3% will lead to fracture. Furthermore, the materials are strong when compressed but weak under tension (16, 17). Another disadvantage of ceramic materials is their vulnerability to flaws or defects in manufacturing related to mechanical, chemical, or thermal processing. Under load, the flaws and the defects can act as a starting point for fracture; a localized concentration of stress will break the atomic bonds and propagate a crack that, if not hindered, will affect the strength of the construction and may cause a complete fracture (18-22).

Ceramic classification In recent years many types of ceramic systems have been introduced in dentistry for all kinds of restoration. Dental ceramics may

21

be classified in a variety of ways based on their intended use, chemical composition, processing method, sintering temperature, microstructure, translucency, fracture resistance, and abrasiveness (23-25). Additionally, the International Organization for Standardization (ISO), ISO 6872:2008 Dentistry – Ceramic Materials (26), lists other classifications: ceramic products that are provided from powder (Type I), and all other forms of ceramic products (Type II) (27). This thesis classifies dental ceramics according to their chemical composition and their intended use (Table 1).

Dental glass-based/porcelain systems Dental glass-based and porcelain systems consist mainly of silicon dioxide (silica or quartz) and varying amounts of alumina (or aluminum oxide). These systems also employ feldspars that are mainly made of aluminosilicates found in nature. They contain various amounts of potassium and sodium. Their main application is for aesthetically important areas, monolithic restorations (veneers), or as veneering with strong optical properties (16). They have disadvantages, such as limited strength, but their strength can be increased significantly by etching and use of a resin bonding (28).

Glass ceramicsThe composition of glass ceramics is similar to that of the porcelain category above. The difference is in the modification process in which varying amounts of different types of crystals are either added or grown in the glass matrix. The primary crystal types available today are leucite, lithium disilicate, and fluorapatite. Dental porcelain and ceramics with high glass content include veneers for monolithic all-ceramic and metal frameworks. They can be used as anterior veneer restorations (e.g., as laminate veneers), in crowns, and in the rest of the mouth as onlays and inlays. Glass ceramic made from fluorapatite and litthium disilicate have been subject of several improvements; from layering glass ceramics to molded pressable ceramics and from the lost wax method to machine-milled ceramics. Those materials are used in full monolithic restorations in any region of the mouth, such as crown, inlays, onlays, veneers and FDPs up to premolar region (29-31).

22

Tabl

e 1.

Cer

amic

mat

eria

ls: c

lass

ifica

tion,

mec

hani

cal p

rope

rties

, exa

mpl

es o

f mat

eria

l sys

tem

s, a

nd fa

bric

atio

n te

chni

ques

Typ

e of

cer

amic

Cla

ssifi

catio

nFl

exur

al

stre

ngth

(M

Pa)*

Frac

ture

toug

hnes

s (M

Pa m

1/2 )

*M

ater

ial s

yste

ms

(exa

mpl

es)

Fabr

icat

ion

tech

niqu

e **

Feld

spat

hic

porc

elai

nPo

rcel

ain

50-1

20~1

.0VI

TA V

M®9/

VITA

BLO

CS®

Sint

erin

g &

hard

mac

hini

ng

Leuc

ite-re

info

rced

gla

ss

cera

mic

Gla

ss12

0-18

0~1

.5IP

S Em

pres

s®Si

nter

ing

& h

eat p

ress

ing

Lithi

um d

isili

cate

-rein

forc

ed

glas

s ce

ram

icC

eram

ic35

0-40

0~3

.0IP

S e.

max

® P

ress

IPS

e.m

ax® C

AD

Hea

t pre

ssin

g

Har

d m

achi

ning

&

crys

talli

zatio

n

Gla

ss-in

filtra

ted

alum

inum

oxid

e

450-

500

3.5-

4.0

In-C

eram

® A

lum

ina

Slip

cas

ting

& g

lass

in

filtra

tion

Gla

ss-in

filtra

ted

zirc

onia

-to

ughe

ned

alum

inum

oxid

e

Hyb

rid

cera

mic

650-

700

4.5-

5.0

In-C

eram

® Z

ircon

iaSo

ft m

achi

ning

& g

lass

in

filtra

tion

Den

sely

sin

tere

d

alum

inum

oxi

de42

0-65

03.

0-4.

0Pr

ocer

a® A

lum

ina

Soft

mac

hini

ng &

sint

erin

g

Y-TZ

PO

xide

cera

mic

900-

1200

6.0-

8.0

Proc

era®

Zirc

onia

D

enzi

r®

Har

d m

achi

ning

* fro

m (1

5) In

: Vul

t von

Ste

yern

P. D

enta

l cer

amic

s in

clin

ical

pra

ctic

e. In

: Niln

er K

, Kar

lsson

S, D

ahl B

L (e

d). A

text

book

of f

ixed

pro

sthod

ontic

s: th

e Sc

andi

navi

an a

ppro

ach.

Sto

ckho

lm: G

othi

a 20

13: 2

05-2

22 *

*(32

) Li R

W e

t al.

Cer

amic

den

tal b

iom

ater

ials

and

CA

D/C

AM

tech

nolo

gy: S

tate

of t

he a

rt J P

rosth

odon

t Res

. 201

4 O

ct;5

8(4)

:208

-216

. (33

) Den

ry, I

et a

l. C

eram

ics

for D

enta

l App

licat

ions

: A R

evie

w. M

ater

ials

2010

, 3, 3

51-3

68.

23

Glass-infiltrated oxide-ceramics (hybride-ceramics) Interpenetrating phase ceramics were introduced in the early 1990s. This system consists of partly-sintered alumina or zirconia cores, reinforced by a glass infiltration technique and subsequently veneered with specially developed porcelain to give the restoration the desired aesthetic properties. Its main use is for single crowns and three-unit FDPs (34, 35). Because this technique is sensitive and its manufacturing time consuming, use of this group of ceramics has decreased while use of lithium disilicate glass ceramic and oxide ceramics has increased (36).

Oxide ceramicsPolycrystalline ceramics, also known as high-strength oxide ceramics, have no glass phase and are acid resistant. The two main oxide ceramics in dentistry are densely sintered aluminum oxide and yttria-stabilized tetragonal zirconia polycrystals (Y-TZP), sometimes described as “ceramic steel”. Due to their outstanding biocompatibility, mechanical properties, and relative translucency, these materials have become more commonly used (37, 38). It could be said that Y-TZP has replaced aluminum oxide as the dominant all-ceramic core material for FDPs. As a substructure material, Y-TZP is used with a veneer material consisting of porcelain or glass ceramics that can be produced using layering or pressable ceramics, or by computer-aided design and computer-aided manufacturing (CAD/CAM). Their range of use is wide: orthodontic brackets, endodontic posts/dowels, veneers, crowns, FDPs, and implant abutments (36, 39-41). New research on stabilized zirconia has developed more translucent forms of the material to achieve more aesthetic monolithic restorations (42, 43). Stabilized Zirconium Dioxide Y-TZP was introduced into dentistry in the 1990s. Pure zirconium-dioxide (ZrO2, zirconia), is a metal oxide that was first identified in 1789 (44). In nature, zirconia is polymorphic, meaning that it displays a different equilibrium (stable) crystal structure at different temperatures with no change in chemistry. It exists in three crystalline phases: above 2370 °C, zirconia forms a cubic solid solution; at intermediate temperatures between 2370 °C and

24

1170 °C, the material transforms into a solid solution of tetragonal microstructure; and below 1170°C, it changes to a monoclinic structure. During cooling, this material undergoes a change in crystal structure; it transforms from a tetragonal to a monoclinic structure. This results in a volume increase in the range of 3–5%. The volume increase induces stress within the material and may initiate crack formation. This makes pure zirconia unsuitable at room temperature for structural or mechanical use. Spontaneous crack formation can occur within the material (32, 44).

It is possible to decrease or eliminate these crystal-structure changes by adding one of several oxides that dissolve within the crystal structure of zirconia. Such oxides include yttrium oxide (Y2O3), cerium oxide (Ce2O3), calcium oxide (CaO), or magnesium oxide (MgO), known as stabilizers. These oxides dissolve within the crystal structure of zirconia and make it possible to densify the material in the tetragonal-phase-range to yield a fine-grained microstructure consisting almost completely of tetragonal grains that is metastable at room temperature (stabilized zirconia) (45-47). The transformation of tetragonal-stabilized zirconia into a monoclinic structure is influenced by temperature, vapor, particle size, the micro- and macrostructure of the material, and the concentration of stabilizing oxides. The most common zirconium dioxide in dentistry is stabilized with 3 mol% yttrium oxide (3Y-TZP). With its mechanical properties such as flexural strength (between 900 and 1200 MPa) and fracture toughness (6–10 MPa/m1/2) it is the dental ceramic considered to have the best mechanical properties (Table 1) (44, 48).

Microcracks, flaws, and defects that inherently grow during the thermal and mechanical processes of manufacturing may significantly influence the fracture resistance of dental ceramics (49). Flaws and defects embedded in the material have the potential to crack under local stress. When a crack occurs, each tetragonal grain has the potential to transform into a monoclinic phase. This causes a local increase in crystal volume of about 3–5%, which inhibits or delays further crack propagation (50, 51). This mechanism is known as transformation toughening; it is a one-way progression that gives the material great potential for stress-bearing and wear-resistant applications (25, 52).

25

The strength and structural stability of Y-TZP can also be affected by exposure to fluids such as saliva, water, and steam sterilization, or via microcracks caused by mechanical treatments such as grinding or airborne-particle abrasion. Metastasized zirconia ceramics may spontaneously transform from tetragonal to monoclinic phase, which is known as low temperature degradation (LTD) (53-55). Fluids such as saliva may penetrate the surface into the material, breaking the bonds between the atoms and leading to an increase in volume that stresses the particles and results in subcritical crack growth (56). Several factors influence the LTD behaviour of stabilized-zirconia ceramics, including grain size distribution; phase composition; stabilizer distribution; and sintering conditions such as sintering time, temperature, and atmosphere (55, 57, 58). Compared with other dental ceramic materials, Y-TZP is more opaque. To achieve a more aesthetic Y-TZP restoration, the core material is covered with veneering porcelain. The most significant sources that hinder the translucency of Y-TZP are the grain boundaries and pores (43). Changes in the amount and size of the crystals and pores, as well as in the sintering process, have made the material increasingly translucent (59-61). Translucent Y-TZP is currently designed for use as full anatomical constructions. These monolithic constructions are characterized by painted stains on the surface, or used with different degrees of cut-back in the labial/buccal areas that are then layered with porcelain to achieve better aesthetics.

Manufacturing Procedure The same set of manufacturers produce almost all zirconia raw materials, but quality varies depending on the type of powder granulometry and the compaction production process used for the discs and blocks. These will determine the final microstructure. There are three ways to compress zirconia powder to increase the density of a ceramic body: uniaxial pressing, cold isostatic pressing (CIP), and hot isostatic pressing (HIP). When using the uniaxial press technique, the powder is pressed in one direction, yielding a green body with low density that is most likely to contain inherent stresses (60, 62, 63). The isostatic methods of compression involve pressure from all directions through a liquid or gaseous medium surrounding the compacted part, yielding a more homogenously packed material with higher density and fewer pores and voids. CIP is conducted at room

26

temperature by one of two techniques: wet bag or dry bag. In the wet bag method, the mold is enclosed with powder, placed in a container filled with liquid, and put under pressure. In the dry bag method, the mold is an integral part of the container and compression uses very little liquid. HIP combines pressing and sintering at increased temperature (1400–1500oC) under high pressure in a gas atmosphere, yielding a dense material of up to 99% density. The diffusion process improves densification, healing voids and pores (63). Previous studies also report that zirconia produced by HIP is less sensitive to LTD (64). The first step of the fabrication process is the manufacture of a block or blanks, which can be made from a huge number of ceramics, using one of these methods (32).

CAD/CAM It is possible to produce blanks and blocks made from stabilized zirconia using a subtractive technique in a CAD/CAM system, either by soft or hard machining. Soft machining includes two different blanks: green stage which is pressed only, or white stage where the powder is pressed and partially sintered. The presintered blanks are easy to shape. The restoration is milled out in a framework enlarged by approximately 20-25%. To achieve the final dimensions and maximal strength, the restoration must be sintered after milling (32, 65, 66). Hard machining involves the milling out of fully sintered blanks that are pressed and completely sintered before milling. Its high hardness requires long milling times and causes rapid wear of the machining tools. Surfaces of dental ceramics machined by hard or soft machining and adjustment grinding show extensive formation of deep defects, microcracking with grain refinement, and deep machining grooves. These factors will decrease the strength of the material (67, 68).

Several CAD/CAM systems are available on the market. The majority use presintered blocks or blanks. The introduction and development of CAD/CAM systems has influenced the production of prosthetic reconstructions, and today this technique is dominant. The introduction of CAD/CAM has made it possible to produce all-ceramic FPDs not only with a higher degree of purity and better aesthetics, but also with a high degree of accuracy (32, 65, 66). In each system, the design is more or less dependent on the design

27

limitations of the CAD software and the milling properties of the CAM system. Many of the design parameters can be individually adjusted to fit each case, while other systems have fixed default settings determined by the manufacturers before milling (69).

Effects of the design of all-ceramic dental restorations and in-vitro test setup FDP core designThe design of crowns and bridges has a great impact on the clinical performance of all-ceramic restorations. Complete fracture through both veneer and core can occur. The design possibilities for an FDP may be limited because of the default settings of CAD/CAM systems and may depend on limitations in the CAD software. In-vitro tests and finite element analyses (FEA) have shown that concentrations of stress occur in the connector of all-ceramic FDPs (46, 70, 71). The dimensions of FDP connectors must be large enough to counteract the concentrations of stress that develop in the framework at the intersection of the pontics. The most important factor seems to be connector diameter, but the radius of curvature at their intersection may also be important (Figure 1) (72, 73). In the CAD software, connector dimensions can be set as either occlusal-gingival height (mm) × buccal-lingual width (mm) or as a cross-section (mm2). In this thesis, all frameworks were designed with connector dimensions set as height × width. Longer spans with several pontics have shown decreasing fracture strength in all-ceramic FDPs (74, 75). Veneer thickness can influence the concentration of stress in an FDP, leading to veneer fractures. To reduce the incidence of veneer fracture or chipping, FDP cores should incorporate anatomical designs, making an even veneer thickness possible (16, 76, 77). In-vitro studies have investigated the fracture strength and fracture mode of stabilized-zirconia FDPs using different shapes and sizes of connectors. A small radius with a design of its own, connected with a straight beam, has been shown to reduce the strength of all-ceramic FDPs (78, 79). However, in-vitro studies have used different testing methods, which affects comparability (80-82).

Several studies have explored the optimal design for all-ceramic stabilized zirconia FDP cores. They have suggested that the app-ropriate shape and dimensions for FDPs should include the following:

28

a minimum core thickness of 0.7 mm (36, 83); an overall smooth, rounded, and anatomically shaped core that allows and supports an even layer of veneering material 0.8–2.0 mm thick (16, 84, 85); connector dimensions of at least 3 mm × 3 mm; and U-shaped gingival embrasure areas, preferably with a radius of at least 0.90 mm (11, 86, 87). This design is widely accepted and is currently considered to be the state-of-the-art design for Y-TZP FDPs.

Figure 1 (a and b). Two connector designs of the gingival embrasure, where rx and ry represent the radius of the curvature of the gingival embrasure: (a) The gingival embrasure has a smooth, round, large radius; (b) The curvature of the width of the gingival embrasure has a smaller radius.

In-vitro test setup for all-ceramic FDPsNew materials on the commercial market need to be tested before clinical use to ensure patient safety. An in-vitro test setup can be made in many different ways and the purpose of in-vitro tests is often to determine the specific properties of a material (88). When testing ceramic materials, the test methods must take into account the range and distribution of forces with consideration for the brittle nature of the material (22, 89). Because tensile strength and fracture toughness can be tested in various ways, several test methods are available

29

depending on the material to be tested. Mechanical tests can be performed with various samples of test specimens including beams, discs, or norm crowns, for example (90-92). When evaluating the material and design of dental reconstructions, their geometry must be considered (82, 93-95). This requires extra effort in designing the framework in an in-vitro study to reproduce the complexity of the supporting systems in the oral cavity that create stress patterns in the FDP (94, 96). In-vitro studies should mimic patient-related factors such as supporting bone, periodontal ligament, abutment teeth with varying root anatomies, and dentin cores of different shapes and quality as nearly as possible (96, 97). Furthermore, in-vitro studies should also consider environmental factors such as moisture, repeated loads, support, and temperature changes, particularly for a complex construction like an FDP (98-100). In-vitro testing can use tooth analogues as abutments. Several tooth analogues have been used in various publications. Examples include cast metals, prefabricated analogues (CAD/CAM), and various kinds of resins (37, 79, 101-103). Extracted human or bovine teeth have also been used (81, 104, 105). The choice of tooth analogues can affect the fracture strength of all-ceramic FDPs. Using rigidly supported tooth analogues may hinder the movement of the abutment teeth. In this case, the load-bearing capacity of the FDPs tested will probably increase and consequently affect the result, giving unrealistic fracture data (96, 97, 106). By comparison, a study design that takes into consideration the supporting tissues, dentin cores, periodontal ligament, and cortical bone may produce decreased load-at-fracture values (96, 97, 105, 107).

Follow up and maintenance of dental ceramic restorations Correct diagnosis and prognosis of the intraoral situation is very important to the clinical outcome of ceramic supraconstructions. Despite the advantages of stabilized zirconia, there are weaknesses associated with the material that need to be improved, especially for their FDP application. Several clinical studies report that fractures occur in the connector of Y-TZP-based FDPs. The most commonly reported problem with Y-TZP FDPs covered by hand-layered veneer is chipping of the veneer (108-112). To overcome these complications associated with chip-off fractures, improvements in the mechanical

30

properties of Y-TZP veneers have achieved comparable mechanical properties to veneers used for metal-ceramic FDPs (86, 113). Several other solutions have been developed, such as the overpressing technique and milling out of the veneer using CAD/CAM technology (97, 114, 115). In the case of chipping failure, this can be repaired using bonding techniques and polishing the surface. Major chipping or connector fractures may require complete replacement of the FDP.

What do we know today? Final remarks The use of stabilized zirconia restorations has increased, but many improvements are still needed. To achieve acceptable long-term clinical survival of stabilized zirconia restorations, we need more knowledge about design and how to increase fracture resistance. Furthermore, there is still no consensus regarding the design and diameter of the connector, nor enough information concerning how different veneering techniques affect fracture strength. Hence, further studies are needed to improve the fracture strength over time of FDPs made from stabilized zirconia.

31

AIMS

The overall aim of this thesis is to evaluate how to improve the design of Y-TZP fixed dental prostheses to increase fracture resistance. The thesis also seeks to evaluate the effect of tooth analogues on fracture strength of Y-TZP.

Specific aims

• To investigate how the fracture strength of 3-unit all-ceramic FDPs made of Y-TZP is dependent on different connector dimensions.

• To evaluate how various radii of curvature in the embrasure area of the connector affect the fracture strength of three-unit all-ceramic FDPs made of Y-TZP.

• To investigate how the choice of material used for supporting tooth analogues and support complexity influences test results relating to fracture strength of FDPs made of the brittle mate-rial Y-TZP.

• To investigate how default settings of two CAD/CAM systems affect the fracture strength and the fracture mode of three-unit and four-unit all-ceramic FDPs made from Y-TZP by compa-ring the radii of curvature in the embrasure area of the con-nector.

• To investigate how FDP length (i.e., number of pontics) affects the fracture strength of Y-TZP.

32

• To investigate and compare the fracture strength and fracture mode in 11 groups of multilayer three-unit all-ceramic Y-TZP FDPs of the kind currently most common with respect to the choice of core material, veneering material, manufacturing technique, design of connectors, and the radius of curvature of FDP cores.

33

MATERIALS AND METHODS

This thesis comprises four in-vitro studies. Table 2 (Table 2.) summarizes the various materials and methods explained in this section. For further details, see the “Materials and Methods” sections of each individual study.

Laboratory ProceduresStudy ISpecimen preparation A master model resembling an upper jaw, with teeth 21 and 23 serving as abutments and tooth 22 missing was made in die stone. Abutments 21 and 23 had a 15° angle of convergence and a cervical preparation design with a 120° chamfer. After scanning this master model with a mechanical scanner, the Procera® Forte (Nobel Biocare, Gothenburg, Sweden), design of the various connector dimensions and the radii of the curvatures of the gingival embrasures were made and divided into groups, (Table 3.) using the scanned data from the CAD software (Procera PCMS, version 1.5, build 75 software). The Procera Production Centre (Nobel Biocare, Stockholm, Sweden) then produced a total of 48 anterior all-ceramic FDPs from Procera® Zirconia bridge material following regular production procedures. The FDPs were supported by end abutments and with one pontic, divided into six groups of eight FDPs each according to their connector dimension and design.

34

Tabl

e 2.

Sum

mar

y of

mat

eria

ls an

d m

etho

ds (*

Radi

us o

f the

cur

vatu

re o

f the

gin

giva

l em

bras

ure)

.

ART

ICLE

III

IIIIV

Type

In

-vitr

oIn

-vitr

oIn

-vitr

oIn

-vitr

o

Spec

imen

s/St

udie

s48

FD

P fra

mew

orks

24 F

DP

fram

ewor

ks16

thre

e-un

it FD

P fra

mew

orks

16 fo

ur-u

nit F

DP

fram

ewor

ks70

FD

P s

plit-

file

fram

ewor

ks40

FD

P s

tate

-of-t

he-a

rt fra

mew

orks

Cor

e/Fr

amew

ork

Mat

eria

l Y-

TZP

Y-TZ

PY-

TZP

Y-TZ

P

Ven

eer

Mat

eria

l -

--

VITA

BLO

CS®

, IPS e

.max

® C

AD

VITA

VM®9,

IPS e

.max

® Zi

rPre

ssIP

S e.

max

® c

eram

Uni

tC

onne

ctor

Radi

us*

Thre

e-un

it FD

Ps2

× 2,

2 ×

3, 3

× 3

mm

0.60

and

0.9

0 m

m

Thre

e-un

it FD

Ps3

× 3

mm

Def

ault

0.60

mm

Thre

e- a

nd fo

ur- u

nit F

DPs

3 ×

3 m

mD

efau

lt

Thre

e-un

it FD

Ps6.

80 ×

2.8

0 an

d 3

× 3

mm

Def

ault

and

0.90

mm

Toot

h an

alog

ues

Dur

aLay

®A

lum

inum

, Pol

ymer

, Dur

aLay

®Po

lym

erPo

lym

er

Dig

itizi

ng d

evic

e

Mec

hani

cal s

cann

erM

echa

nica

l sca

nner

Mec

hani

cal s

cann

er an

d Las

er

disp

lace

men

t gau

ge sc

anne

rTw

o di

ffere

nt o

ptic

al sc

anne

rs

Art

ifici

al A

ging

H

eat t

reat

men

t, Pr

eloa

ding

(1

0,00

0 cy

cles

, 30

N to

30

0 N

) The

rmoc

yclin

g (5

000

cycl

es 5

°C to

55°

C)

Hea

t tre

atm

ent,

Prel

oadi

ng

(10,

000

cycl

es, 3

0 N

to

300

N) T

herm

ocyc

ling

(500

0 cy

cles

5°C

to 5

5°C

)

Hea

t tre

atm

ent,

Prel

oadi

ng

(10,

000

cycl

es, 3

0 N

to

300

N) T

herm

ocyc

ling

(500

0 cy

cles

5°C

to 5

5°C

)

Hea

t tre

atm

ent,

Prel

oadi

ng

(10,

000

cycl

es, 3

0 N

to

300

N) T

herm

ocyc

ling

(500

0 cy

cles

5°C

to 5

5°C

)

35

Table 3. Study I: Connector diameter and radius of the curvature of the gingival embrasure (GE radius).

Group Height Width GE radius

1 2 mm 2 mm 0.6 mm2 2 mm 2 mm 0.9 mm3 3 mm 2 mm 0.6 mm4 3 mm 2 mm 0.9 mm5 3 mm 3 mm 0.6 mm6 3 mm 3 mm 0.9 mm

Heat treatmentAll FDP cores underwent heat treatment in a calibrated porcelain furnace (Multimat Touch & Press; Degudent, Dreieich, Germany) to simulate the firing cycles of the veneering porcelain (Nobel Rondo™ Zirconia, Nobel Biocare) manufacture recommended veneering porcelains for the core material used. Heat treatment comprised four firing programs: Liner, Dentin 1, Dentin 2, and Glaze, all according to the manufacturer’s instructions.

Supporting tooth analoguesThe fabrication of supporting tooth analogues was created by reproducing the abutments of the master cast, first and third incisor, were reproduced using A-silicon impression material (President, Coltene AG, Altstätten, Switzerland), to make two tooth-like end abutments on which the FDPs could be cemented and fixated in acrylic blocks. Impressions were taken of the central incisor and canine abutments using A-silicon impression material (President, Coltene AG) then poured in die stone (Vell-mix, Kerr, Romulus, MI, USA) with a metal dowel pin centred in each abutment to stabilise the following build-up of a wax-up. Grooves were made on the dowel pins to facilitate retention of the wax. The above-mentioned reproduction abutments were to be used only to create the shape for the final test model and were not to be used in the test. They were copied in a second step using an A-silicone impression (President, Coltene AG) and subsequently poured with inlay pattern resin (DuraLay®, Reliance Dental MFG Co., Worth, Illinois, USA) which was weighed and measured according the manufacture recommendation, creating the final model for the experiment (Figure 2). A total of 48 tooth analogues, 96 copies, 48 incisors and 48 canines were made.

36

Figure 2 (a-d). Fabrication of supporting tooth analogues: (a) First and third incisor reproduced from master impression poured in die stone, (b) Wax-up of rots, (c) Mold for duplicating wax-up, (d) Incisor and canine in DuraLay®.

CementationThe DuraLay® tooth-supporting analogues were cleaned and treated with ED primer II A and B (Kuraray Medical INC.) which was applied to the cementation surfaces according to manufacturer’s instructions.

The FDPs of all groups were luted onto the reproduction abutments with Panavia F 2.0 luting cement (Kuraray Medical INC.) using both Oxyguard II (Kuraray Medical INC.) and light curing lamp (Ivoclar-Vivadent Bluephase, Scaan, Liechtenstein). Polymerization light inten-sity was 1100 mW/cm2 and curing time was 20 seconds in each of four directions, 90° apart, and then 60 seconds in one direction with the seating load removed, according to the manufacturer’s recommendations. During setting of the cement, all FDPs were loaded in the direction of insertion with a force of 15 N for a period of 60 seconds. The cemented FDPs were fixated in holes in acrylic blocks with die stone (Vell-mix, Kerr, Romulus, MI, USA) subsequent to cementation before starting of ageing of the FDPs cores. To create moist environment of the oral cavity, the FDPs were placed in a plastic container, with a sealable lid,

37

and with distilled water (37±1°C). The FDP cores remained in this moist environment throughout the testing period until load-to-fracture.

Artificial aging - thermocycling and cyclic preload Artificial aging comprised two steps. The first stage of aging – thermocycling (TC) – put each FPD through 5000 cycles in two water baths at temperatures of +5°C and +55°C, respectively. Each cycle lasted 60 seconds: 20 seconds in each bath and 10 seconds to complete the transfer between baths. In the second stage of aging, all specimens underwent cyclic preloading of 30 N and 300 N in a wet environment and mounted at a 10° inclination relative to the vertical plane. Cyclic preloading comprised 10,000 cycles with a load profile in the form of a sinusoid wave at 1 Hz. A stainless steel indenter, 2.5 mm in diameter, applied the force centrally on the incisal edge of the lateral incisor pontic to avoid sliding during loading (Figure 3).

Figure 3 (a-d). A schematic picture of the jig and the application of the load during preloading and load-to-fracture: (a) brass foundation, (b) tooth analogue, (c) the pontic, (d) stainless steel indenter.

Load to fractureThe specimens were placed in a testing jig at a 10° inclination and subjected to a load applied by a universal testing machine (Instron

38

4465, Instron Co. Ltd, Norwood, MA, USA). The crosshead speed was 0.255 mm/min and the load was applied with a stainless steel intender 2.5 mm in diameter, placed centrally on the incisal edge of the lateral incisor pontic. Fracture was defined as a visible fracture through the entire construction.

Statistics Means of the Student’s t-test determined the differences in fracture strength between the groups at a level of significance set to p ≤ 0.01.

Study IIPreparation of master casts for three-unit FDPsFor the 3-unit mater cast, plastic model of an upper jaw (KaVo Dental GmbH, Biberach, Germany) was used. Two abutment preparations were made on the 21 incisor and the 23 canine with missing teeth 22. The aim was to design a preparation with a 120° chamfer and an angle of convergence of 15°. Subsequently a full arch A-silicone (Flexitime Mono Phase, Heraeus Kulzer, GmbH, Hanau, Germany) impression was taken and poured with die stone (Everest® Rock, Type 4 die stone, KaVo Dental) to produce a master cast (Figure 4).

Figure 4 (a-d). Master model production: (a-b) Plastic model with two abut-ment preparations, (c) Impression of plastic model, (d) Master cast in die stone.

39

Specimen preparation The master cast was scanned once with a mechanical scanner, Procera® Forte (Nobel Biocare, Zurich, Switzerland). The connector dimensions were set to 3 mm x 3 mm and the minimum thickness of the core was set to 0.7 mm. The radius of the gingival embrasure in the connector areas was 0.6 mm in accordance with both the default settings of the CAD program and the manufacturer’s recommendations. A total of twenty-four FDPs in Procera® Zirconia bridges were produced following standard production methods.

Supporting tooth analoguesStudy II required three types of tooth analogue (Figure 5). Two of the tooth analogues were milled using the same CAD file as the one used to produce the FDP cores at the production center (Procera® Production center, Stockholm, Sweden), and the other, complex tooth analogue was made using the same technique described in Study I. The milled tooth analogues were made from either aluminum or a polymer material (polyoxymethylene-copolymer [POM-C]). The complex tooth analogues were made of DuraLay® (Reliance Dental MFG Co., Worth, Illinois, USA). A total of 24 tooth analogues divided in 3 groups were to be made.

Figure 5 (a-c). Tooth analogues: (a) Tooth analogues milled from aluminum, (b) Tooth analogues milled from polymer material, (c) Complex tooth ana-logues made in DuraLay® cemented to a Y-TZP FDP, fixated in acrylic block with die stone.

Heat treatment and artificial ageing - TC All specimens received heat treatment using the same firing program as Study I and in the same manner. Furthermore, all FDP cores underwent TC (LTC Multifunctional Thermocycler; LAM Technologies Electronic Equipment, Firenze, Italy) for 5000 cycles in the same manner as Study I using a small basket controlled by a device driver.

40

CementationThe tooth analogues made from aluminum received steam cleaning before treatment with a metal adhesive primer (Alloy primer; Kuraray Medical Inc., Okayama, Japan). The tooth analogues made from polymer and those made from DuraLay® were steam cleaned and subsequently treated with ED primer II A and B (Kuraray Medical Inc.). The specimens were luted onto the reproduction abutments with Panavia F 2.0 luting cement (Kuraray Medical INC.) using both light and Oxyguard II (Kuraray Medical INC.), all according to the manufacturer’s recommendations. All specimens cemented to the tooth analogues in Study II used the same cementation procedures as in Study I and were stored in the same manner as in Study I, both to create a moist environment like the oral cavity and to prevent desiccation of the luting cement.

Cyclic preload and Load to fractureAll specimens received cyclic preloading of 10,000 cycles and were loaded to fracture in the same manner as Study I.

Statistics Means of the Student’s t-test determined differences in fracture strength between the groups with the level of significance set to p ≤ 0.001. Fisher’s exact probability test calculated differences in fracture modes at a level of significance of p ≤ 0.01.

Study IIIPreparation of the master cast for three-unit and four-unit FDPs The master cast of the three-unit was the same master cast used in Study II. The mater cast of the four-unit were made with the same design and material as the 3-unit mater cast. Two abutment preparations were made of the right central incisor and the left canine, the left central incisor and the left lateral incisor were removed (Figure 6).

41

Figure 6. A plastic model with two abutment preparations, incisor 11 and canine 23, for preparation of the four-unit master cast.

Specimen preparation Study III used two different scanners to produce the specimens for four groups with two different designs and units of two different lengths. The master casts were scanned with a mechanical scanner, Procera® Forte (Nobel Biocare, Zurich, Switzerland) and the data was transferred to a computer equipped with CAD software (Procera CAD Design C3D, version 2.60) where the intended design of the FDPs was established. The connector dimensions were set to 3 mm x 3 mm and the minimum thickness of the core was set to 0.7 mm. The radius of the gingival embrasure in the connector areas was 0.6 mm according to the default settings in the CAD program and according to the manufacturer’s recommendations. To generate data for the last two groups, the master cast was scanned with a laser displacement gauge scanner, NobelProcera® Scanner (Nobel Biocare, Zurich, Switzerland). The data was transferred to a computer equipped with CAD software (NobelProcera 3D prosthetic design, version 4.1.10.4.). This software produced the intended design of the FDPs according to the protocol established for the first two groups. A total of 32, all-ceramic Y-TZP FDP cores: 16 three-unit and 16 four-unit cores with differing design were made. These FDPs had either one or two intermediate pontics, supported by end abutments. The FDPs were then divided according to design in four groups, with eight FDPs in each group.

Heat treatment and artificial aging - thermocycling All specimens received heat treatment and TC for 5000 cycles in the same manner as Study II.

42

Supporting tooth analoguesThe production center (Procera® Production Centre, Stockholm, Sweden) produced the tooth analogues for the testing procedure at the same time it produced the FDPs from the same CAD file as the one used to produce the Y-TZP FDP cores, as the same manner as in Study II. A total of 32 tooth analogues, 16 three-unit and 16 four-unit inspection blocks, from a polymer material.

CementationThe tooth analogues received steam cleaning and subsequent treatment with ED primer II A and B (Kuraray Medical Inc.). The specimens were luted onto the reproduction abutments with Panavia F 2.0 luting cement (Kuraray Medical INC.) using both light-cured with a curing lamp (Heraeus Translux® Power Blue®, Hereaus Kulzer GmbH, Hanau, Germany) and Oxyguard II (Kuraray Medical INC.), all according to the manufacturer’s recommendations. All specimens cemented to the tooth analogues were cemented and stored in the same manner as Studies I and II, to create a moist environment of the oral cavity and to prevent desiccation of the luting cement.

Cyclic preload and load-to-fractureAll specimens underwent cyclic preloading for 10,000 cycles and load-to-fracture in the same manner as Studies I and II.

StatisticsMeans of the Student’s t-test determined differences in fracture strength between the groups with the level of significance set to p ≤ 0.05.

Study IV Preparation of the master cast In this study, the master cast of the three-unit FDP was the same master cast used in Study II and Study III.

Specimen preparation To prepare the specimens in Study IV, two different optical scanners were used. Data for Groups 1-7 (the split-file core design) were generated using the Sirona InEos Blue (Sirona Dental Systems GmbH, Germany). The master cast was scanned once and the

43

data was transferred to a computer equipped with CAD software (Sirona inLab, version 3.88) which produced the intended design of the FDP. Data for Groups 8-11 (the state-of-the-art core design) were generated with the 3shape D640 (3Shape A/S, Copenhagen, Denmark). The master cast was scanned once and the data was transferred to a computer equipped with CAD software (Dental-designer 3shape 2013, build 2.8.8.0) which produced the intended design of the FDP. A total of 110 anterior three-unit Y-TZP FDP cores with one intermediate pontic, supported by end abutments, were made (Figure 7), (Table 4.). In addition to the 110 FDPs, we made two extra cores, one of each design for analyzing the connector cross-section areas.

Figure 7. Overview of all-ceramic FDPs, Groups 1–11: core design, core material, and veneer techniques.

Split-file design cores The split-file design was made in total of 70 FDP split-file cores, divided into seven groups. In four groups the FDPs were made in VITA In-Ceram® YZ for inLab®, YZ-40/15 (VITA Zahnfabrik,

44

Bad Säckingen, Germany). In three groups the FDPs were made in IPS e.max® ZirCAD for inLab MO 0 B40 (Ivoclar Vivadent AG, Liechtenstein). The connector dimensions of the Y-TZP cores in the split-file groups were 5.40 mm × 2.50 mm for the left central incisor and 6.80 mm × 2.80 mm for the left canine, with a bar-shaped occlusal design in accordance with the default settings of the CAD program. The core thickness was 0.70 mm. The radius of the curvature of the gingival embrasure in the connector areas was selected according to the default settings of the CAD program and the manufacturer’s recommendations (Figure 8).

Figure 8 (a-b). (a) Split-file core design: Circle shows the radius of the gingival embrasure; (b): Enlargement of the radius of the curvature of the gingival embrasure.

Veneering of the split-file core The aim was to create a split-file design structure that allowed a veneering material with a thickness of 1.5 mm. The CAD data for the FDPs were sent to a milling machine Sirona inLab MCXL (Sirona Dental Systems GmbH, Germany), which produced the split-file designed FDPs. The design of the veneering materials in the four milling groups followed the manufacturer’s recommendations. In one of the groups, the material used for the milled veneer structures was prepared from VITABLOCS® (VITA Zahnfabrik) and subsequently luted onto the substructure with Panavia F 2.0 luting cement (Kuraray Medical Inc.). In two of the groups, the material used for veneer milling was IPS e.max® CAD (Ivoclar Vivadent AG), which was subsequently fused to the substructure with IPS e.max® Crystall./

45

Connect (Ivoclar Vivadent AG). The veneers for the overpressing group were milled from combustible acrylic blocks of IPS AcrylCAD® (Ivoclar Vivadent AG), then mounted onto the cores, and finished with over-pressing according to the lost-wax method using IPS e.max®

ZirPress (Ivoclar Vivadent AG). All frameworks of the multilayer veneer were attached following the manufacturers’ instructions.

State-of-the-art core design The state-of-the-art cores comprised four groups, containing a total of 40 FDPs. Two of the groups were made in VITA In-Ceram® YZ DISC, Ø 98 x 18 mm (VITA Zahnfabrik). The remaining two groups were made in BruxZir® HT 2.0, Ø 98 x 15 mm (Glidewell Laboratories, California, USA). In the CAD program, the design of the connector dimensions of the Y-TZP were set to 3 mm × 3 mm, the minimum thickness of the core to 0.7 mm, and the radius of the gingival embrasures in the connector areas to 0.90 mm. The CAD data for the FDPs were subsequently sent to a milling machine Wieland 4030MN (Wieland Dental, Technik GmbH, Germany), with the software CAM: Zenotec CAM 2.2.017, (Wieland Dental, Technik GmbH, Germany) where they were used for production of the FDPs (Figure 9).

Figure 9. State-of-the-art core design with connector dimensions 3 × 3 mm, radius of the curvature of the gingival embrasure 0.90 mm.

46

Tabl

e 4.

FD

P sp

ecim

ens

in S

tudy

IV. C

onne

ctor

hei

ght a

nd w

idth

(H/W

), ra

dius

of t

he c

urva

ture

of t

he g

ingi

val e

mbr

asur

e, n

umbe

r of

uni

ts, m

ater

ial,

and

core

des

ign.

Gro

upC

ore

Y-TZ

PV

enee

rH

/Wα

H/W

βRa

dius

γU

nit

Cor

e de

sign

1VI

TA In

-Cer

am®

CO

RE*

5.40

× 2

.50

mm

6.80

× 2

.80

mm

defa

ult

3-un

itSp

lit-fi

le

2VI

TA In

-Cer

am®

VITA

BLO

CS®

5.40

× 2

.50

mm

6.80

× 2

.80

mm

defa

ult

3-un

itSp

lit-fi

le

3VI

TA In

-Cer

am®

IPS

e.m

ax®

CA

D5.

40 ×

2.5

0 m

m6.

80 ×

2.8

0 m

mde

faul

t3-

unit

Split

-file

4VI

TA In

-Cer

am®

VITA

VM

®9

5.40

× 2

.50

mm

6.80

× 2

.80

mm

defa

ult

3-un

itSp

lit-fi

le

5IP

S e.m

ax® Zi

rCA

DIP

S e.

max

® Z

irPre

ss5.

40 ×

2.5

0 m

m6.

80 ×

2.8

0 m

mde

faul

t3-

unit

Split

-file

6

IPS e

.max

® Zi

rCA

DIP

S e.

max

® C

AD

5.40

× 2

.50

mm

6.80

× 2

.80

mm

defa

ult

3-un

itSp

lit-fi

le

7IP

S e.m

ax® Zi

rCA

DIP

S e.

max

® c

eram

5.40

× 2

.50

mm

6.80

× 2

.80

mm

defa

ult

3-un

itSp

lit-fi

le

8VI

TA In

-Cer

am®

CO

RE**

3 ×

3 m

m3

× 3

mm

0.90

mm

3-un

itSt

ate-

of-th

e-ar

t

9Br

uxZi

r® H

T 2.

0C

ORE

**3

× 3

mm

3 ×

3 m

m0.

90 m

m3-

unit

Stat

e-of

-the-

art

10VI

TA In

-Cer

am®

VITA

VM

®9

3 ×

3 m

m3

× 3

mm

0.90

mm

3-un

itSt

ate-

of-th

e-ar

t

11Br

uxZi

r® H

T 2.

0VI

TA V

M®9

3 ×

3 m

m3

× 3

mm

0.90

mm

3-un

itSt

ate-

of-th

e-ar

t

α C

onne

ctor

dim

ensi

on in

the

left

cent

ral i

ncis

or. β

Con

nect

or d

imen

sion

in th

e le

ft ca

nine

. γ R

adiu

s of g

ingi

val e

mbr

asur

. *C

ore

only

– te

sted

with

out v

enee

r mat

eria

l. Sa

me

core

des

ign

as g

roup

1-7

. **C

ore

only

– te

sted

with

out v

enee

r mat

eria

l. Sa

me

core

des

ign

as g

roup

8-1

1

47

Manually built-up porcelain To duplicate the shape of the veneers in the milled veneering groups one of the completed IPS e.max® CAD-on FDPs was placed on a tooth analogue and the undercuts were waxed. Then duplicated the shape of the FDP using A-silicon and putty impression material (PRESIDENT, Coltène AG, Altstätten, Switzerland). The performed manual layering of the veneering porcelain was made in four groups, two spilt-file and two with state-of-the-art core design. To separate the layered porcelain from the inner surfaces of the impression, we used a GI-MASK® universal separator for silicones, Coltène® (Coltène/Whaldent AG, Switzerland) and thereafter applied LPC Isolating Liquid (Ivoclar Vivadent AG). The VITAVM®9 (VITA Zahnfabrik, Germany) porcelain, was used to veneer two groups of state-of-the-art cores and one of the two split-file core groups. The last group of split-file cores received IPS e.max® Ceram (Ivoclar Vivadent AG, Liechtenstein). Veneering ceramic for dentin was applied on the FDPs frameworks using the impression to achieve standardized shape and size of the FDPs. All the layered FDPs then underwent a first firing. A second layer dentin ceramic was applied to compensate the sintering shrinkage (Figure 10) and finally applied glaze to the FDPs. All porcelain firing proceeded in a calibrated porcelain furnace (Ivoclar P 500; Ivoclar Vivadent AG) according to the firing programs (Table 5-6) recommended by the manufacturer’s instructions.

Heat treatment and artificial aging - thermocycling Three groups comprised core material only. One group had a split-file core made of VITA In-Ceram® YZ (VITA Zahnfabrik). The other two groups had state-of-the-art cores design with one group made of VITA In-Ceram® YZ (VITA Zahnfabrik) and one of BruxZir® HT 2.0 (Glidewell Laboratories). The FDP cores in the three groups were sub-jected to heat treatment to simulate the firing cycles of the veneering porcelain VITAVM®9 (VITA Zahnfabrik, Germany) and IPS e.max®

ceram (Ivoclar Vivadent AG, Liechtenstein) according to the manu-factures recommendations. The group FDPs made of VITABLOCS® (VITA Zahnfabrik, Germany), which was subsequently luted onto the substructure with Panavia F 2.0 luting cement (Kuraray Medical INC) were excluded from the heat treatment procedure according to the manufacturer’s recommendations. All specimens received thermo-cycling for 5000 cycles in the same manner as Studies II and III.

48

Figure 10 (a-d). Manually hand-layered built-up porcelain: (a) IPS e.max® CAD-on FDP placed on tooth analogue, undercuts waxed, (b) Mold for duplicating FDP silicon material, (c) Veneering layering split-file core, (d) first firing of layered FDP.

Supporting tooth analoguesThe tooth analogues for the testing procedure were made from the same CAD file as the one used in Studies II and III at the production center (Procera® Production Centre). A total of 110 three-unit inspection blocks of the polymer material were made.

CementationThe tooth analogues received steam cleaning and subsequent treatment with ED primer II A and B (Kuraray Medical Inc.). The specimens were luted onto the reproduction abutments with Panavia F 2.0 luting cement (Kuraray Medical Inc.) using both light-curing with a curing lamp (Heraeus Translux® Power Blue®, Heraeus Kulzer GmbH) and Oxyguard II (Kuraray Medical Inc.), all according to the manufacturer’s recommendations. All specimens cemented to the tooth analogues were cemented and stored in the same manner as Studies I-III to mimic the moist environment of the oral cavity and prevent desiccation of the luting cement.

49

Cyclic preload and load-to-fractureBefore load-to-fracture, all specimens underwent cyclic preloading for 10,000 cycles using a specially constructed preloading device (MTI Engineering AB, Lund, Sweden/Pamaco AB, Malmö, Sweden). Cyclic preloading and loading-to-fracture proceeded in the same manner as in Studies I-III.

AnalysisAfter load to fracture, all presented studies in the present thesis were examined and analysed, both visually and under a light microscope (Leica DFC 420, Leica Application Suite v. 3.3.1, Leica Microsystems CMS GmbH, Wetzlar, Germany) to establish the fracture modes. Moreover, two FDPs in Study IV, one a state-of-the-art and one a split-file core, were cut in the connector areas with a diamond saw (IsoMet® 5000 Liner Precsion Saw, BUEHLER, Lake Bluff, USA) and measured the cross-sectional areas under the light microscope.

Statistical Analysis The Student’s t-test determined any differences between groups with the significance set to p ≤ 0.05.

50

Tabl

e 5.

Gen

eral

Firi

ng P

rogr

am IP

S e.

max

® C

eram

.

Firi

ng P

rogr

amPr

ehea

ting

tem

p.Pr

ehea

ting

Dry

ing

time

(min

)H

eatin

g Ra

te

Y-TZ

P

Firi

ng te

mp.

Hol

ding

tim

e

(min

)

Vac

uum

1V

acuu

m 2

ZirL

iner

firin

g40

3 °C

450

°C

960

°C1

450

°C71

9 °C

Was

h Fou

ndat

ion

403

°C4

40 °

C75

0 °C

145

0 °C

749

°C

Den

tin fi

ring

140

3 °C

440

°C

750

°C1

450

°C74

9 °C

Den

tin fi

ring

240

3 °C

440

°C

750

°C1

450

°C74

9 °C

Gla

ze fi

ring

403

°C6

60 °

C72

5 °C

145

0 °C

724

°C

Tabl

e 6.

Gen

eral

Firi

ng P

rogr

am V

ITAVM

®9.

Firi

ng P

rogr

amPr

ehea

ting

tem

p.Pr

ehea

ting

Dry

ing

time

(min

)H

eatin

g Ra

te

Y-TZ

P

Firi

ng te

mp.

Hol

ding

tim

e

(min

)

Vac

uum

(min

)

Was

hbak

e firi

ng50

0 °C

255

°C

950

°C1

8.11

Den

tin fi

ring

150

0 °C

655

°C

910

°C1

7.27

Den

tin fi

ring

250

0 °C

655

°C

900

°C1

7.16

Gla

ze fir

ing

with

VI

TA A

KZEN

T®

500

°C4

80 °

C90

0 °C

1-

51

RESULTS

Study I Fracture strength of three-unit fixed partial denture cores (Y-TZP) with different connector dimension and design.

All the FDPs fractured in the area of the connector. The crack propagation that led to fracture started at the gingival side of the connector. Two of the FDPs with connector dimensions 2 × 2 mm and radius of 0.60 mm lost retention at one abutment during thermocycling (No. 2 and 6). They were re-cemented and following the remaining preparation method. Table 7 provides the fracture data and statistics.

All the 48 FDPs in the present study were examined to establish the fracture modes (Figure 11) were distributed as follows:

(A) In 30 FDPs the fractures began in the corner between the pontic and the connector and extended through the pontic.

(B) In eight FDPs, the fractures began in the connector area at the corner between the pontic and the connector and extended through the connector.

(C) In five FDPs, the fractures began in the middle of the connector beam and extended through the pontic.

(D) In five FDPs, the fractures began in the corner between the pontic and the connector and extended through the middle of the connector.

52

Six FDPs fractured during the cyclic mechanical preload (30-300 N): three with connector dimensions of 2 × 2 mm and a radius of 0.60 mm (No. 2, 5, and 7), and three with connector dimensions of 2 × 2 mm and a radius of 0.90 mm (No. 1, 3, and 7).

Table 7. Results of Study I: Fracture load (N), mean fracture load, and standard deviation. The groups represent different connector dimensions (2 × 2, 3 × 2, and 3 × 3 mm) and different radii of curvature (0.60 and 0.90 mm). There were significant differences between FDPs of the same dimensions but different radii.

Group 1 2 3 4 5 6

Dimension Radius

2×2 mm 0.6 mm

22 mm 0.9 mm

3×2 mm 0.6 mm

3×2 mm 0.9 mm

3×3 mm 0.6 mm

3×3 mm 0.9 mm

Significance NS NS p ≤ 0.01

FDP No.1 442 300 β 512 580 755 730

2 300 α’ β 455 673 726 489 800

3 353 300 β 633 581 670 987

4 388 375 615 521 580 873

5 300 β 499 621 554 499 783

6 349 α 498 426 460 803 739

7 300 β 300 β 669 590 659 839

8 434 469 482 599 754 899

Mean(N) 358 399 579 576 651 831

SD 58 91 93 76 119 87α FDPs which lost retention at one abutmentβ FDPs fractured at 30-300 N preloadNS: No significant difference

53

Figure 11. Crack propagation directions in Study I.

Study II The influence of support properties and complexity on fracture strength and fracture mode of all-ceramic fixed dental prostheses.

Table 8 lists the fracture data from Study II. All the FDPs fractured in the connector area, which corresponds with previous studies of all-ceramic FDPs. Fracture modes differed significantly (p ≤ 0.01) between the aluminum group and the polymer and DuraLay® groups, which were distributed as follows:

• In Group A (aluminum) 75% of the FDPs fractured through both connectors and 25% fractured through one connector only (Figure 12).

• All the FDPs in Group P (polymer) fractured through one con-nector only (Figure 13) (six of the FDPs fractured through the mesial part of the connector and the pontic, and two fractured in a similar manner but through the distal part of the connec-tor and the pontic).

• All the FDPs in Group D (DuraLay®) fractured through the distal part of the connector and through the pontic.

54

Figure 12. FDP fractured through both connectors in the Group A aluminum tooth analogues in Study II.

Figure 13. FDP fractured through the distal part of the connector and the pontic in the Group P polymer tooth analogues in Study II.

All crack propagations that led to fracture began at the gingival side of the connector. None of the abutments failed during testing. Fracture loads were significantly higher in Group A compared to Group P (p ≤ 0.001) and Group D (p ≤ 0.001). There was no significant difference in fracture loads between Group P and Group D.

55

Table 8. Results of Study II. Load at Fracture (N).

Load at fracture (N) Aluminum (A) Polymer (P) DuraLay® (D)

FDP core no. 3-unit Y-TZP 3-unit Y-TZP 3-unit Y-TZP

1 1562 753 700

2 1704 749 568

3 1728 849 675

4 2103 735 647

5 1646 767 724

6 2189 830 728

7 1964 610 829

8 1641 575 875

Mean 1817 733.5 718

SD 235 96 98

Significance p ≤ 0.001 NS NS

NS: No significant difference

Study IIIFracture strength of all-ceramic (Y-TZP) three- and four-unit fixed dental prostheses with different connector design and production history.

All FDPs fractured in the area of the connector, with crack propagation leading to fracture. Table 9 lists the fracture data. During cyclic mechanical preload (30-300 N), four of the FDPs in group 4Z:1 fractured.

All the 32 FDPs in the present study were examined to establish the fracture modes, which were distributed as follows:

• In group 3Z:1, six FDPs fractured at the distal connector area on the left central incisor and through the pontic. Two FDPs fractured at the mesial connector area of the left canine and th-rough the pontic. All FDPs fractured at the sharp mesial corner of the connector.

56

• In group 4Z:1 seven of the FDPs fractured at the mesial area of the connector on the left lateral incisor. One FDP fractured at the mesial connector area of the left canine and through the pontic. All FDPs fractured at the sharp mesial corner of the connector (Figure 14 a).

Figure 14 (a-b). Four-unit FDP designs in Study III: (a) Fracture mode Group 4Z:1, (b) Fracture mode Group 4Z:2.

• In group 3Z:2 four of the FDPs fractured at the distal connec-tor area on the left central incisor and through the pontic. Four FDPs fractured at the mesial connector area on the left canine and through the pontic. All fractures were located in the center of the connector area where the connector is thinnest.

• All the FDPs in group 4Z:2 fractured at the mesial connector area on the left lateral incisor and through the pontic. All frac-tures were located in the center of the connector area where the connector is thinnest (Figure 14 b).

A significant difference was found between Groups 3Z:1 and 3Z:2 (p ≤ 0.05), and Groups 4Z:1 and 4Z:2 (p ≤ 0.05).

Study IVInfluence of core design, production technique and material selection on fracture behavior of Y-TZP FDPs produced using different multi-layer techniques: split-file, over-pressing and manually built-up veneers.

57

Table 9. Results of Study III: fracture load (N), mean fracture load, and standard deviation.

Group 3Z:1 3-unit 3Z:2 3-unit 4Z:1 4-unit 4Z:2 4-unit

Significance p ≤ 0.05 p ≤ 0.05

FDP No.1 753 805 300* 5752 749 953 300* 5123 849 697 499 4554 735 1070 300* 5645 767 843 508 4306 830 997 530 5747 610 921 300* 5038 575 999 499 547

Mean(N) 734 910 405 520

SD 96 122 112 55

*FDPs fractured at 30-300 N preload

Table 10 lists the fracture data. The fracture mode was determined by three different failure types: cohesive (chipping of the veneering ceramic), radial cohesive veneer fractures, or total fracture through the whole construction. The fracture modes were distributed as follows:

• All the FDPs in Group 1 fractured completely in the central connector, starting gingivally with a crack growth towards the loading site on the pontic or involving the connector only.

• In Group 2, seven of the FDPs fractured as described above for group 1. Three FDPs showed radial cohesive veneer fractures: one FDP fractured in the disto-buccal connector area of the pontic, and two FDPs fractured in the mesio-buccal connector area of the pontic (Figure 15).

• In Group 3, seven of the FDPs fractured as described above for Group 1. Three FDPs showed radial cohesive veneer cracks: one FDP in the mesio-buccal connector area of the pontic and two FDPs in both the mesio-buccal and the disto-buccal connector/pontic areas.

58

• All the FDPs in Group 4 fractured in the lateral connector, starting gingivally with a crack growth towards the loading site on the pontic or involving the connector only.

• The FDPs in Group 5 fractured in either the central connector (n=6) or the lateral connector (n=4), all starting gingivally with a crack growth towards the loading site on the pontic or invol-ving the connector only.

• The FDPs in Group 6 fractured in either the central connector (n=6) or the lateral connector (n=2), all starting gingivally with a crack growth towards the loading site on the pontic or in-volving the connector only. Two FDPs showed radial cohesive veneer cracks in the mesio-buccal connector area of the pontic.

• All the FDPs in Group 7 fractured as described above for Group 1.

• In all the FDPs in Groups 1-7, the fracture was initiated at the sharp indentation in the gingival embrasure area.

• The FDPs in Group 8-11 fractured in either the central connec-tor or the lateral connector, all starting gingivally with a crack growth towards the loading site on the pontic or involving the connector only. The distribution was as follows:

• Group 8, Three fractures in the central connector and seven in the lateral one.

• Group 9, Two fractures in the central connector and eight in the lateral one.

• Group 10, All FDPs fractured in the lateral connector.

• Group 11, Two fractures in the central connector and eight in the lateral one.

In groups 8-11, all fractures were located in the center of the connector area where the connector dimension is thinnest. None of the 110 FDPs had chip-off fractures.

59

Figure 15. Visible radial cohesive veneer fracture in Group 2.

60

Tabl

e 10

. Re

sults

of S

tudy

IV: g

roup

s 1-

11, l

oad-

at-fr

actu

re (N

)

Load

at

G-1

G-2

G-3

G-4

G-5

G-6

G-7

G-8

G-9

G10

G11

frac

ture

(N)

Split

-file

des

ign

Stat

e-of

-the-

art d

esig

n

FDP

core

no.

185

818

2014

6418

2016

3615

5015

7511

4411

0717

1115

24

210

7714

75*

1660

1539

1860

1424

*14

2111

7010

2516

3817

09

399

919

2717

08*

1748

1908

1488

1314

1154

1103

1801

1519

498

219

0116

57*

1399

1785

1917

1604

1037

1117

2162

1754

511

2816

5414

0919

0917

8414

6014

5811

6510

5018

0916

15

690

417

3813

1417

1718

0917

18*

1536

1084

945

1909

1658

710

4720

5413

77*

1730

1906

1597

1596

1113

1206

1919

1795

811

1517

9716

3414

8318

5813

7415

9311

2210

2419

8215

74

995

619

44*

1550

1573

1907

1703

1467

1093

1172

1785

1695

1099

717

46*

1702

1741

2085

1507

1593

1235

1084

1774

1559

Mea

n10

0618

0615

4816

6618

5415

7415

1611

3210

8318

4916

40

SD88

165

145

160

115

164

9755

7615

097

* FD

Ps w

ith ra

dial

coh

esiv

e ve

neer

frac

ture

und

er lo

ad to

frac

ture

61

Statistical AnalysesTable 11 presents the statistical analysis.

Table 11. Statistical analysis for Study IV (p ≤ 0.05).

Split-file design State-of-the-art design

G1 G2 G3 G4 G5 G6 G7 G8 G9 G10 G11

G1 * * * - - - * - - -G2 * * NS - - - - - - -G3 * * NS - NS - - - - -G4 * NS NS - - - - - * -G5 - - - - * * - - - -G6 - - NS - * NS - - - -G7 - - - - * NS - - - -G8 * - - - - - - NS * -G9 - - - - - - - NS - *G10 - - - * - - - * - *G11 - - - - - - - - * *

* = Significant difference- = Not comparable, or irrelevant NS = No significant difference

62

DISCUSSION

Methods: In-vitro studies and test setup When a new material is released, researchers often undertake in-vitro studies to evaluate the material and design in an attempt to prevent and minimize clinical failure. Compared to in-vivo studies, in-vitro studies are relatively efficient in terms of time and cost, but possess several limitations. The test is simplified and does not encompass the complexity of the biological environment of the oral cavity and the loads of chewing. Therefore, drawing parallels between in-vitro studies to predict clinical results may prove difficult, and the results from in-vitro studies should be interpreted with care. However, the test setup needs to be performed in as clinically relevant a manner as possible and to take many aspects into consideration. These tests can be performed in many different ways and should be selected according to the needs of a controlled laboratory environment (88, 96, 97).

Choice of specimen designThe use of all-ceramic material has increased in dentistry, especially zirconium dioxide-based restorations. When clinical failures occur, they can be related to the shape and geometry of restorations. Understanding how to mimic the environment in the oral cavity is very important in laboratory tests, as mentioned above (86, 113).

In Studies I-III the test specimens were frameworks without a layer of veneering material. It was assumed that the presence of veneering material would not influence the results significantly and therefore excluded the veneer. Another reason for this exclusion was that

63

applying veneer manually can never achieve an identical thickness between specimens, which would likely influence the results. The purpose of Study I was to evaluate the effect of increasing connector dimensions and the radius of the curvature of the gingival embrasure on the fracture strength and fracture mode of FDPs. Study II sought to evaluate how different tooth analogues affect the fracture strength and fracture mode of FDPs tested with identically designed frameworks. Study III evaluated the different framework default settings of two different CAD/CAM systems. It also compared how different radii of curvature in the embrasure area of the connector affect the fracture strength and fracture mode of three-unit and four-unit constructions with number of pontics was made. Study IV evaluated the fracture strength and fracture mode of two substructure designs, including three veneer techniques with different types of veneer producers. This study used two types of connector designs with various dimensions and radii of curvature in the gingival embrasure. Three different manufacturers produced the framework material for Study IV.

CAD/CAM systemOver the last decades, the use of CAD/CAM applications in dentistry has steadily increased. Most likely even more systems will be released in the future. CAD/CAM in dentistry has made it possible to produce prosthetic restorations much faster, more cheaply, and more precisely, and has enabled the use of new materials such as oxide ceramics. The systems vary depending on the needs of construction and material to be milled (65, 116). Dental CAD/CAM systems can be assembled using several types of unit, for example, a scanner connected to milling centers (production of inlay to long-span of FDP), in-house systems (complete systems, production from inlay to long-span of FDP) or chairside systems (inlay to short-span of FDP) (116, 117). A variety of scanners are currently available to capture tooth preparation in 3D. Such scanners can be mechanically based with a contact probe, or a digital scanner laser displacement gauge, or a line laser beam with a camera (116, 117). Unfortunately, several scanners have proprietary systems, forcing the operator to send the scan file to one specific manufacturer. Studies I and II used a mechanical scanner to scan the master models and then sent the scan files to a milling center. In the FDP designs in Study I, the CAD software allowed changes in the radius of the gingival and occlusal embrasure from 0.60 mm to

64

0.90 mm. In Study II the CAD software had been updated and the parameters to increase the radius were removed and replaced with a default setting of 0.60 mm. Study III used two scanners, a mechanical scanner and a scanner laser displacement gauge. Designing the FDPs was done in default setting according to the manufacturers recommendations. The FDPs produced by the mechanical scanner had small radius in the connector area with square-shape design. The FDPs produced by the laser scanner had a larger, U-shaped radius with a smooth and rounded design Figure.16. Study IV used two different in-house CAD/CAM systems, both using optical scanners. In one of the systems, the FDP’s core and veneer were milled and joined together. This kind of FDP will be bulky. The design of the connector and the radius were 5.40 mm × 2.50 mm in the left central incisor and 6.80 mm × 2.80 mm in the left canine with a bar-shaped occlusal design in accordance with the default settings in the CAD program. The second CAD/CAM system was an open system allowing the operator to use any brand of material. The design of the connector dimension used a state-of-the-art design with a connector of 3 mm × 3 mm and a U-shaped radius of gingival embrasure of 0.90 mm. The properties of CAD/CAM systems do not only depend on the CAD software. CAM devices can also mill in several axes in various materials. The number and capability of the CAM’s axes and the use of different sizes and shapes of diamond and cutting tools greatly impact the final outcome of dental restorations (50, 118).

Figure 16. Design of the connector dimensions of a Y-TZP four-unit core with default U-shaped radius design, scanned with a laser displacement gauge.

65

Veneer Techniques Study IV was the only study in this thesis that used veneer materials. It evaluated three techniques of veneer production: milling, pressing, and manual layering. To achieve decreased numbers of clinical veneer fractures several techniques has been developed with an aim to overcome human performance inconsistencies, as well as to improve quality, reliability and cost-effectiveness. The CAD/CAM system that was able to produce split-file FDPs produced the milling and overpressing groups. The milled resin replicas for overpressing technology served as molds in which the veneering ceramic was heated and pressed into the investment mold directly over the Y-TZP framework (97, 119). Comparing the veneer group luted onto the substructure with groups joined to the core with low-fusing ceramic e.max® CAD-on Crystall./Connect was rather technique sensitive, because it required vibration. The use of two groups veneer connected with e.max® CAD on Crystall./Connect was made to encompass evaluations of substructures. The default veneer designs in the CAD/CAM system with the split-file framework had a veneer thickness of 1.5 mm. The manually built-up porcelain and ceramic groups were duplicated using one completely milled veneer and core joint together. Using translucent stabilized zirconia as a substructure allowed evaluation of the material with and without veneer. The aim was to create identical veneer thicknesses between FDPs. Manually applied veneers can never be made identical with one another. The liquid and powder used in combination with its compaction will inevitably influence veneer thickness and may affect the results.

Supporting tooth analoguesThe FDPs in Study I were cemented onto abutments made of inlay pattern resin. Additionally, in Study II, one of three groups was cemented in this same manner. The two other groups of tooth analogues in Study II were made by CAD/CAM of polymer polyoxymethylene-copolymer (POM-C) and titanium at the same time as production of the FDPs and were produced from the same CAD file. Furthermore, the tooth analogues made in Studies III and IV on three-unit FDPs was made of this same polymer material, reusing the same CAD file as in Study II. The four-unit FDPs cemented on tooth analogues in Study III were made of the same

66

material and in the same manner as the Study II polymer group. The inlay-pattern-resin group was complex, being made in several steps and cemented with various components in order to mimic the clinical situation. Even following standardized methods, human error will affect accuracy if tooth analogues are handmade. Tooth analogues made by CAD/CAM are precisely milled in one unit. Choosing the appropriate tooth analogues to test ceramic crowns and bridges is a topic that deserves further scrutiny. The choice of material for the tooth analogues could significantly influence results. Rigid tooth analogues and rigid supports that hinder the movement of the abutment teeth will increase the load-bearing capacity of the FDPs in in-vitro studies (96, 97, 114). Consequently, this will affect results, giving unrealistic fracture data (96, 97, 114). Nevertheless, using rigid tooth analogues to simulate and evaluate implant abutments is more relevant than tooth analogues tested with FDPs (97). Likewise, comparative studies that considered supporting tissues, dentin cores, the periodontal ligament, and cortical bone found decreased load-at-fracture values (105, 120). In the clinical situation, crowns and bridges are supported by a combination of different structures with varying properties, such as their E-modulus, and this must be considered in in-vitro studies. Selected the tooth analogues in Studies I-IV were made after reviewing the material’s E-modulus (Table.12). The results of Study II provided the basis for the use of tooth analogues in Studies III-IV.

Cementation and water storageAll studies used the same resin cement: 10-methacryloyloxydecyl dihydrogen phosphate. Cementation procedures were the same in all studies and standardized according to the recommendations of the manufacturer. It received an applied load of 15 N during polymerization in an apparatus, light-curing, and an application of Oxyguard II. Furthermore, after cementation all frameworks were stored in distilled water at 37±1°C. By covering the bottom surface and sealing the lid to create a humid atmosphere, aimed to prevent desiccation of the luting cement. The FDPs were stored in this manner all through the artificial aging processes until the load-to-fracture test.

67

Table.12. Examples of the E-modulus of structure materials.

Example of Structures E-Modulus (GPa)*

Enamel ≈ 75-120Dentin ≈ 18-25

Pulp ≈ 0.0002

Cortical bone ≈ 14

Cancellous bone ≈ 0.25

Periodontal membrane ≈ 12-15

DuraLay® ≈ 2.6

Polyoxymethylene-copolymer POM-C ≈ 2.8

Acrylic resin ≈ 8.3

Artificial teeth ≈ 8.3

Plaster Type IV ≈ 15

Plexiglas ≈ 15

Porcelain ≈ 100

Zirconia ≈ 200

Aluminium ≈ 70

Titanium ≈ 103-120

* from: (121) WJ O’Brien. Dental material and their selection: Quintessence Publishing Co,Inc: Chandler Drive 2008: 330-335. (122) IA Ibrahim. Particulate reinforced metal matrix composites - a review: Journal of materials science 26.5 (1991): 1137-1156. (16) Anusavice,K.J.; A textbook Dental Ceramics, Phillips’ science of dental materials. 12th ed. St. Louis, Saunders 2013 418-473. (123) Nordbergs tekniska AB.

Artificial aging (heat treatment, thermocycling, mechanical preload)The aging of crown and frameworks made of all-ceramic materials in the oral environment may influence their durability. An attempt to mimic the ongoing aging of oral restorations was made. The in-vitro studies incorporated several artificial aging processes to decrease the strength of the material and bond. However, there is no consensus regarding the best test setup procedures for aging (36, 98, 124). All frameworks tested in Studies I-IV underwent the same aging procedure: heat treatment, TC, and mechanical preload. The FDPs in all four studies, except one group of FDPs with luted veneers in Study IV, underwent heat treatment, i.e. porcelain firing, overpressing, or crystal fusion. When a stabilized zirconia core is veneered, it is subjected to firing at high temperatures (750–950oC) and subsequent

68

cooling. Usually, two to five firing cycles are required to obtain an aesthetically acceptable restoration. The temperatures that the core experiences during porcelain firing may decrease the mechanical properties of the ceramic material, an effect called low-temperature degradation (LTD) (53, 125). Furthermore, dental restorations are subjected to a moist environment and different temperatures. It is possible to test dental restorations using TC, but the extent to which TC mimics the oral environment has been questioned. There is no consensus regarding number of cycles, temperatures, dwell time, and intervals between baths (98). Exposure of stabilized zirconia to the oral environment with saliva or other fluids may start a progression of material degradation (53). The TC temperature used in this thesis of an approximate 50oC difference in temperature between the two baths may not affect the ceramics, but rather the enclosed materials. Several other investigations have used TC limits of the range 5oC to 55oC (81, 98, 120, 126). In addition, 10,000 thermal cycles correspond to approximately one year of clinical function (98). In Study I, the frameworks were cemented with the tooth analogues and then exposed to TC. Two tooth analogues abutments came loose during the TC, the cement was removed and were recemented, then subjected to the same treatment as the other specimens. In Studies II-IV, all frameworks FDPs underwent TC before cementation.

To simulate oral chewing, cyclic preload can be combined with water to evaluate flexural and fracture strength. The numbers of cycles and magnitude of load can vary between studies (36, 80, 127). All the frameworks used in this thesis received a load fluctuating between 30 N and 300 N for 10,000 cycles in distilled water. The presence of water is recommended in preload conditions to closely simulate clinical use (127). Other laboratory studies may apply fewer cycle loads but many more cycles (99, 128). In study IV, additional supplement were made to prevent clinically irrelevant Hertzian cone cracks of the veneer materials, by using a thin plastic foil during the cyclic preloading (129, 130).

Load-to-fracture All the FDPs in this thesis underwent load until fracture. The test was performed in water with force applied centrally with a steel indenter on the incisal edge of the pontic. In Study IV, a 1 mm thick plastic

69

foil was placed between the steel intender and centrally on the incisal edge of the pontic to prevent Hertzian cone cracks. Studies I-III used no such foil. The test setup may, under controlled conditions, provide basic but valuable information, such as the fracture strength of the different materials. Hence, the test method depends on the material to be tested, taking into account the specific properties of each material. Tests that can measure the tensile strength of a material include uniaxial tensile strength tests; three- or four-point bending tests; and biaxial bending tests, which include the ring-on-ring, ball-on-ring, and piston-on-three-balls tests (131-133). Fracture toughness can be measured using various modifications of indentation tests (134, 135). Tests of dental materials are performed using standardized stylized beams, discs, or norm crowns (90, 133, 136, 137). The outcome of the in-vitro test is usually complete fracture through core and veneer. However, this differs from results in clinical reports in which the most common complication is fracture of the veneer of stabilized zirconia (41, 113, 138). In Study IV, all groups with milled veneers had more than two FDPs that did not experience complete fracture through the core and veneer.

Discussion of results Intraoral loads and mechanical preloadEstimations of the clinical loads dental restorations experience during normal function and maximal bite force vary among publications (139-141). The number of chewing cycles per day is estimated at approximately 800–1400 (142). Maximal bite forces in the oral cavity have a wide span that ranges from a few hundred Newtons in the anterior region to over 1000 Newtons in the posterior region. Men achieve higher forces than women by maintaining a higher mechanical force for the muscles of mastication (141). It can be assumed that during normal mastication, the bite forces are considerably lower than the maximal bite forces registered in the anterior region (143). Different factors such as gender, age, dysfunction, and denture use lower the reported values (139, 140, 144). A dental ceramic restoration needs to withstand maximum and normal repeated forces over time. Furthermore, it is necessary they have safety margins to withstand complex clinical forces that cannot be completely reproduced in an in-vitro test. However, several FDP frameworks in Studies I and III fractured during mechanical

70

preloading at a load of 300 N or below. It could not be established at what preload value below 300 N the fracture occurred; therefore all frameworks fractured values were recorded to 300 N. Comparing the results, it seemed possible that the data from these fracture values below 300 N would produce a significant difference but, as the data were inaccessible, it could not be confirm.

FDP design The connector/radius The design of connector dimensions is of great importance in an FDP. When total fracture of FDPs occurs clinically, the fracture is located in the connector (108-110, 145). These findings are corresponding to the results in this thesis. However, there is no consensus regarding the design of the connector or its dimensions for in-vivo and in-vitro studies. A previous clinical study indicated that the connector’s cross-sectional area was under-dimensioned when fracture occurred (74). The cross-sectional area of the connector is set in CAD software via vertical height and horizontal width or is reported in mm2. Several CAD software programs have default settings with mm2 that may limit optimal connector design. Previous clinical and laboratory studies indicated that the height of the connector may be more important than the width in achieving durable all-ceramic FDPs (100, 146). The results of Study I agree with these results. Moreover, the results in Study I indicate that fracture strength may increase with an optimal connector design with margin of dimension and an enlarged U-shaped radius of the gingival embrasure. Our results agree with published in-vitro and finite element analysis (FEA) studies (11, 46, 49, 87, 147, 148). Study IV shows that, designs without an enlarged U-shaped radius of the gingival embrasure of connector, as in a state-of-the-art design (3 mm × 3 mm, radius 0.90 mm), as compared to split-file design, require a connector with at least a 42% larger cross-sectional area in order to achieve the same load-to-fracture strength as the state-of-the art design. Two framework designs made of Y-TZP were used. The first framework connector design (split-file) was 5.40 mm × 2.50 mm in the left central incisor and 6.80 mm × 2.80 mm in the left canine with a bar-shaped occlusal design in accordance with the default settings of the CAD program. The second framework, a state-of-the-art design, was 3 mm × 3 mm with a U-shaped radius of 0.90 mm. As measured under a light microscope (Leica DFC 420),

71

the cross-sectional areas were 11.2 mm2 (Figure 17) and 17.9 mm,2 respectively, for the two connectors in a split-file core design.

Figure 17. Measurement of a cut split-file Y-TZP FDP core under a light micro-scope: the figure shows the cross-sectional area of the connector between the central incisor and pontic.

The frameworks made according to the state-of-the-art design had a cross-section of 7.3 mm2 (Figure 18). The split-file designed cores required a connector with a cross-sectional area nearly 42% larger in the vertical aspect (or 62% if comparing with the largest connector), to achieve nearly the same fracture values as the state-of-the-art cores. An increased radius at the gingival embrasure also increases the fracture load, whereas the radius of curvature at the occlusal embrasure only has minor effect (11, 73). Recommendations for connector dimension and radius of the gingival embrasure of Y-TZP FDPs in study I was: anteriorly minimum 3 mm occlusal-gingival in height and 2 mm in buccal-lingual width with 0.90 mm radius. The connector design of 3 mm × 3 mm with a radius of 0.90 mm withstood fracture loads above 800 N. Such a design should be considered for zirconia-based FDPs with longer spans or molar replacements. Clinical studies show that three-unit zirconia-based FDPs perform well (9, 109, 111) but survival rates decrease with increasing numbers of pontics (74, 146). Study III evaluated three-unit and four-unit FDPs with connector dimensions of 3 mm × 3 mm and various radius designs.

72

Figure 18. Measurement of a cut state-of-the-art Y-TZP FDP core under a light microscope: the figure shows the cross-sectional area of the connector between the canine and pontic.

Using two different scanner systems with default settings, the laser scanner systems produced a connector design more in agreement with the state-of-the-art design. Comparing the two types of connector design by measuring the radius of the curvature of the gingival embrasure under a light microscope (Leica DFC 420), the default settings of the laser scanner produced a radius that was approximately 20% larger (Figure 19-20). In addition, the groups that were produced using the laser scanner systems achieved 20% higher fracture strength. Furthermore, clinical findings investigating longer FDP spans (74, 138) and the results of Study III show the need for increasing connector dimensions with proper radius design. When increasing a three-unit to a four-unit FDP, the fracture strength decreases by approximately 50%. Moreover, a previous laboratory study showed a decrease in fit between the restoration and abutment tooth with increased span length in stabilized zirconia FDPs (149). The results of Studies I, III, and IV indicate that when the height of the connector increases in combination with a small radius of the gingival embrasure, fracture strength is less than with a state-of-the-art design connector of 3 mm × 3 mm and a radius of 0.90 mm.

73

Figure 19. Four-unit FDP in Study III (group 4Z.1) under a light microscope: the figure shows the measured radius of the curvatures of the gingival embrasure (circles diameter).

Figure 20. Four-unit FDP in Study III (group 4Z:2) under a light microscope: the figure shows the measured radius of the gingival embrasure (circle diameter).

74

The state-of-the-art framework design in Study IV withstood fracture loads above 1100 N. In comparison with results from Studies I and III of three-unit FDPs with similar connector designs, fracture load was higher in Study IV. This could be related to the manufacturing process and use of CAD/CAM system. The core thickness was set to 0.70 mm in all studies in this thesis. A previously published clinical study (150) evaluated the fracture of three-unit stabilized zirconia FDPs with fractography analysis. The reason for fracture was due to under-dimensioned core thickness, which was set to 0.30 mm (150). Furthermore, a previously published in-vitro study indicated that core thickness may be reduced in an all-ceramic Y-TZP FDP without reducing its strength (151).

Milling dental ceramics The use of CAD/CAM machining has become much more common for milling all-ceramic restorations. Despite the excellent mechanical properties of stabilized zirconia, these FDPs may fail after a short period of time, in particular due to veneer chipping (86, 109). Although stabilized zirconia ceramics are well-known for phase transformation toughening, this becomes less effective at the weak parts of the restorations, like the connectors and the cervical margin where micro-defects and residual stresses are concentrated. The entire production process of zirconia restorations creates microdefects and residual stresses that are generated over several steps such as blank production; CAD/CAM-milling and sintering; and clinical adjustment by dental technicians or dentists through grinding, abrasion, and polishing (91, 118, 152). Though sintering can relieve the stresses introduced by milling, the microdefects remain (60, 153). A final evaluation that was done after the tests were performed in Studies I-III found that the rough pattern and stripes of the FDP cores occurred during the milling process as a result of the shape of the diamond or milling tools (Figure 21). Under a light microscope (Leica DFC 420), accumulated micro-defects could be observed (Figure 22). These microdefects created residual stresses and acted as fracture indentations. By comparison, this rough pattern or these stripes in the FDPs milled by the two separate CAD/CAD systems in Study IV was not found. This may be one reason why the three-unit FDPs with a state-of-the-art connector design (3 mm x 3 mm, radius 0.90 mm) in Study IV exhibited higher fracture strengths than the FDPs in Study I.

75

Figure 21. Rough pattern and stripes after milling in a four-unit FDP from Study III (group 4Z:2).

Figure 22. Surface of a four-unit FDP in Study III (group 4Z:2) under a light microscope: the figure shows microdefects.

LTDFracture strength may have decreased due to the aging processes that all stabilized zirconia FDPs in this thesis underwent. The aging processes may have created the problem of sensitivity to LTD (126, 154, 155). LTD occurs within the temperature range of 65–500oC

76

and is related to a surface transformation from the metastable tetragonal-phase (t) to the stable monoclinic-phase (m), occurring especially in the presence of water or water vapor (57, 156). The mechanism on the surface leads to degradation, which propagates into the bulk material through micro- and macrocracking. This weakens the mechanical properties (53, 54). Microdefects arise during milling, airborne-particle abrasion, or grinding of stabilized zirconia material. This makes the material more vulnerable and more exposed to LTD. Repeated loading in combination with the presence of water increases the mechanism for the t → m transformation. Transformation may create cracks along the grain boundaries that, in turn, allow moisture to penetrate further into the material (99, 125). Oral environmental factors, such as constant humidity and repeated occlusal loads, may accelerate the aging process and reduce the mechanical properties of the materials over time (126). It is hard to predict LTD in a clinical situation over time. Actual clinical conditions may be even worse than predicted due to the additional effect of stresses.

Veneering techniques and fractures It is well known that veneer fractures are the most common clinical complication with porcelain veneer fused to stabilized zirconia FDP cores. Study IV examined three veneer techniques for an all-ceramic FDP: conventional hand-layering, overpressing, and milling. None of the veneered FDPs had cohesive chips, which is the most commonly reported clinical failure (78, 86, 108, 109, 111, 112, 138, 157). Instead, fractures occurred as visible radial cohesive cracks (Figure 23) (158, 159). All other FDPs fractured in the connector area, which is in agreement with Studies I-III. In all veneer groups, there was an obvious and statistically significant increase in fracture strength compared to the non-veneered groups, which shows that the complete material systems are more reliable than each component on its own (114). Hand-layered veneer ceramic demonstrated the lowest mean fracture strength value, 1516 N, while the overpressed ceramics demonstrated the highest mean fracture strength value, 1854 N. In recently published in-vivo studies (72, 160) overpressed ceramics showed fewer veneer fractures, while built-up hand-layered veneer showed more veneer fractures. All veneered FDPs in Study IV

77

showed results that exceeded the expected average maximum loads with a large safety margin, indicating sufficient fracture strength. One reason for veneer fracturing may be the milling technique. Two of the groups were joined with ceramic IPS e.max® CAD to the Y-TZP core with low-fusing ceramic Crystall./Connect materials, whereas the third porcelain VITABLOCS® group was luted with resin cement, which demonstrated significantly higher fracture strength than did the two other groups (158). In a previously published FEA study comparing conventional veneer to stabilized zirconia and veneer luted to stabilized zirconia under load, less stress developed in the veneer luted to the stabilized zirconia (161). Another possible explanation for the radial cohesive cracks could be that the milling process creates microdefects leading to crack initiators on the inner surface of the veneering material (118, 159). Furthermore, the intermediate luting and crystal fusing may not adequately and homogeneously fill the defects created by the milling process (114, 158, 162).

Figure 23. Cut split-file FDP (Group 2 in Study IV) under a light microscope: the figure shows a radial cohesive crack.

78

A third possible explanation could be that the process for fusing the veneer groups with Crystall./Connect to the core is a technique-sensitive method (158). The risk of porosities is probably higher when attempting to cover a relatively large surface, as in Study IV where the FDPs consisted of a three-unit construction rather than the single units tested in other studies (Figure 24) (114, 163, 164). Study IV shows reversed results compared to other in-vitro studies of the CAD-on technique. Their findings show that the CAD-on technique creates fewer or no radial cohesive veneer fractures due to the homogeneity of the veneering material (114, 163, 164). However, studies made on single constructions that are load-bearing and tested on tooth analogues made from stiff metal are not comparable to three-unit FDPs tested on polymer tooth analogues as in Study IV (114, 163, 164).

Figure 24. Cut CAD-on FDP (Group 3 in Study IV) under a light microscope: the figure shows porosities in the Crystall./Connect surface with arrows.

Four veneered groups used the porcelain build-up layering technique. Two of these had split-file design substructures made from two different material systems, but had exactly the same design. The group layered with conventional veneering porcelain was compared to a layer with a veneering nano-flourapatite glass-ceramic. There was a significant difference between the two groups, favoring

79

porcelain veneer, which corresponds with the results of previous studies (165, 166). Only one of the two groups was made with highly translucent stabilized zirconia whereas both groups were made with substructures designed according to state-of-the-art design and layered with conventional porcelain. The porcelain veneer fused to Y-TZP with the state-of-the-art substructure design presented the highest mean fracture strength values of 1849 N. Difference in stress distribution between a single crown and an FDP with connectors and pontics may also explain the differing results (167). Besides issues related to manufacturing processes, porosities, and micro defects, it is still unclear why veneer failures occur. Various studies have evaluated the reasons for material fractures (36, 156, 168) suggesting many possible reasons.

Veneer failure predictions Multifaceted factors The highest numbers of complications from the use of stabilized zirconia in prosthodontic treatments occurs with FDPs. Studies have reported a variety of factors that may influence chipping. Laboratory tests, such as FEA, may help to predict fracture behaviour of specific material combinations. Clinical failures may be mainly influenced by preparation angle and finishing line design, internal fit and cement thickness, or the individual FDP design (86, 156). The veneer fractures of stabilized zirconia (chipping, cracking, delaminations, and large fractures) are higher than the veneer fractures observed for metal-ceramic restorations (36, 109, 110). The literature has discussed many reasons for failure of veneered stabilized zirconia, including intraoral load, FDP design, grinding, ceramic microdefects, aging and LTD, and veneer techniques, all discussed in the text above. Furthermore, thermal stress, inappropriate veneer thickness, defects in the veneering material, and multifaceted bond failures have all been suspected to result in veneer fractures.

Thermal stresses In every firing cycle when veneering stabilized zirconia, the physical properties change from visco-elastic to a solid state. This is followed by a change in volume and, consequently, in density. Changes in length/volume that usually depend on temperature and heating/

80

cooling rate occur when cooling down from sintering to room temperature (63). Residual thermal stresses throughout the veneer layer may occur as a function of the rate of cooling (169). Inside the veneer, the residual stress could be produced by coefficient thermal expansion mismatches, by the temperature gradients produced during cooling, or both (170, 171). In the temperature gradient, which is the difference between the inner and outer temperatures of the FDP surface, high tensile stresses will occur on the surface when it reaches room temperature (170, 172). In the firing cycle, the cooling process leads the outer FDP surface to become solid first, and it starts to contract, inducing a compressive stress. Simultaneously the inner surface is in a visco-elastic state at a high temperature. Moreover, the thermal conductivity between stabilized zirconia and the veneer may increase the intensity of the thermal gradient within the veneering porcelain (172, 173). The faster the cooling, the higher the glass transition temperature (Tg), which cause stresses to develop (156, 172). With slow cooling by delaying the furnace door opening, a decrease in residual thermal stresses within the veneer layer has been observed (49, 174-176). In all studies in this thesis, the firings proceeded according to the manufacturer’s recommendations. Most manufacturers of porcelain veneers for Y-TZP restorations have modified their firing guidelines and adopted a slow-cooling-rate protocol.

Inappropriate veneering thickness Support of the proper veneering ceramic and production of even thickness are factors that need to be considered from the start of treatment planning. Clinically evaluated veneer chippings have frequently been associated with lack of appropriate porcelain support (74, 109, 113, 177). To achieve clinically acceptable aesthetics, manufacturers recommend the veneering thickness chosen in Study IV. Most manufacturers recommend guidelines on minimal ceramic thickness ranging from 1.5 mm to 2 mm (178, 179). Suggestions of what constitutes appropriate shape and dimensions have been made and these include a minimum thickness and an overall smooth and rounded anatomically shaped core allowing and supporting an even layer of veneering material 0.8–2.0 mm thick (16, 180). All FDP framework veneers in Study IV were designed to have an even,

81

1.5-mm veneer thickness. The results of fracture strength in the hand-layered veneer groups showed a significant difference between the two framework design groups, favoring the state-of-the-art group. Veneer thicknesses of 1 mm or thinner in the hand-layered group might have increased this difference in the results of Study IV (180). Thickness recommendations are mostly based on the outcome of in-vitro tests (178).

Defects in the veneering material Porcelain fracture is frequently related to internal and subsurface flaws. This will affect the strength of dental ceramics. Clinical analyses clarify that flaws are one of the causes of veneer-zirconia fractures (170, 181). In most cases, fractures initiated by flaws inside the material result in defects in marginal areas or at the occlusal surface (118, 153, 170). The flaws within milling veneer ceramics originate from the manufacturing process of the blanks. In hand-layered porcelain/ceramics, the defects originate from the porcelain powder, dental laboratory handling, or contamination in clinical use when fitting and adjusting the restoration (67, 156, 168).

Multifaceted bond failureMany reasons for veneer fractures have been discussed. Furthermore, veneer fractures are most probably multifaceted in origin, in particular arising from bond failures (156, 168, 181). Daou et al listed several potential causes related to bond failure. These include differences in thermal coefficients and liner materials, and poor core wetting; veneer firing shrinkage; phase transformation; loading stresses; flaw formation; coloring pigments; and surface properties (168). None of the studies in this thesis used coloring pigment or colored Y-TZP.

Choice of tooth analogues In-vitro studies that evaluate all-ceramic crowns, FDPs of three-units or longer spans, and implant restorations and which test mechanical loads, fracture strength, and fracture mode can choose from an assortment of possible tooth analogues. An evaluation of other studies testing three-unit, all-ceramic FDPs on tooth analogues found no obvious correlation between load-at-fracture and choice of material for the tooth analogues (96, 97).

82

Studies using stiff tooth analogues to test FDPs showed unrealistic fracture strengths compared to clinical studies. However, many other factors affecting fracture load may exist when testing all-ceramic FDPs (67, 156, 168). Moreover, varying the design of FDPs can test for differences between FDPs placed anteriorly or posteriorly (137).

The height and cervical shape of the preparation may affect the result, as can the amount of root in the tooth analogue mounted in the test block if there is one (96, 97, 181). A short, wide preparation mounted with a large part of its supported root in the test block will probably have a higher load-at-fracture, while a tall, narrow preparation with a smaller part of the supported root in the test block will presumably have a lower one. Study II, the FDPs were all identical and exposed to the same test setup. Results for the stiff tooth analogues made from aluminum showed 250% higher loads-at-fracture than the two other tested tooth analogues. The outcome of in-vitro studies testing all-ceramic FDP dental material with stiff tooth analogues may be misleading concerning choice of material for clinical use. For clinically relevant strength, the fracture strength of a material intended for cementation should range between 300 N and 800 N, while the two other tested tooth analogues in Study II showed load-at-fracture values of over 700 N. In tests of all-ceramic materials by in-vitro study, the main aim should not be to achieve excessively high results, rather, it should be to mimic natural teeth. However, every step in the test setup of a study must be clearly presented or it calls into question the significance of that study. The use of stiff tooth analogues is relevant when evaluating all-ceramic materials for implant FDPs (97). Fracture mode analyse Analysis of the fracture modes in Studies I-IV found no direct connection between the modes. However, from manufacturing the FDPs until load-at-fracture, there are many different aspects that may affect the fracture mode: blank production, flaws, milling defects, tooth analogues, products, etc. Studies I and II used complex handmade tooth analogues, and manual skills may have affected the cementation of the preparation-rots resin to the Plexiglas®. However, the E-modulus of the stiff tooth analogues in Study II influenced

83

the fracture mode (97). The fracture modes in Study III evaluating four-unit FDPs of Y-TZP indicate that smoother connectors and larger radii of the curvature of the gingival embrasure will lead to a crack penetrating from the gingival embrasure of the connector into and through the pontic against the applied load-to-fracture. This observation is apparent in all the FDP designs with smoother connectors and larger radii of the gingival embrasure, as well as with state-of-the-art designs. Moreover, the milling tooth analogues made with polymer in Studies II-IV had higher root anatomy on the left lateral incisor, which corresponds to most fracture modes. However, the reason for fracture modes depends on a variety of factors.

Clinical significance In recent decades, stabilized zirconia restorations have become increasingly accepted, and the material is now one of most common materials used in dentistry. Several clinical trials have been done on stabilized zirconia FDP restorations, but more long-term follow-up is needed (67, 86, 113). Many improvements have been made, but many challenges to understanding the clinical complications of stabilized zirconia restorations remain. To minimize clinical complications, such as veneer chipping and complete fracture of stabilized zirconia FDPs, clinicians and dental technicians need to understand the material’s sensitivity and design properties. The state-of-the-art design gives approximately 25% higher fracture strength. Furthermore, it is very important that dental laboratories evaluate the properties of CAD/CAM systems when they invest in new equipment. CAD/CAM systems should be open and CAD software should allow for designing FDPs with varying connector dimensions and radii of the gingival embrasure. The clinician needs to try out the FDP core before veneer applications to make sure the dental laboratory has delivered proper design. This gives an opportunity to achieve proper veneer thickness, fitting, hygienic space, and so on for a better all-ceramic stabilized zirconia FDP.

Future investigations The past 30 years of development of biomedical engineering in dentistry have made the evolution of dental ceramics interesting. Introduction of stabilized zirconia ceramics has opened a wide range

84

of all-ceramic applications. Considering the clinical complications of veneering material fractures, careful planning of further in-vivo studies should give attention to the FDP core design. Veneering materials and different veneer techniques have sufficiently improved for all-ceramic cores, and this will play a significant role in their long-term success. However, to overcome clinical complications, evolution towards monolithic materials has been introduced and further clinical studies are needed.

Remarkable progress has been made in ceramic processing and development. In dentistry, new technical breakthroughs with additive manufacturing in biomedical engineering are on the horizon, and this will open up possibilities for unlimited geometries and applications.

85

CONCLUSIONS

The recommended minimum connector dimension of an anterior three-unit Y-TZP FDP is 3 mm in the incisal-cervical direction and 2 mm in the buccal-lingual direction.

Increasing the radius of the gingival embrasure from 0.6 mm to 0.9 mm, will increase the fracture strength of a Y-TZP FPD with a connector dimension of 3 mm × 3 mm by 20%.

There is a need for a simple, standardized, test setup to achieve realistic and comparable results when in-vitro testing all-ceramic FDPs.

Both resilient non-complex and resilient complex tooth analogues gave comparable test results when the test setup was unchanged in all other aspects.

Non-resilient (with an E-modulus of aluminum or higher) tooth analogues gave high and unrealistic load-at-fracture values together with adverse fracture modes compared to FDPs involved in clinical failure.

The default settings of the two different CAD/CAM systems had a great impact on fracture strength. It is important that a CAD/CAM system is equipped with the ability to design a connector that fulfils the clinical demands of mechanical function and longevity.

The design and the size of the radius of the gingival embrasure is crucial to load-bearing capacity, but also affects the fracture mode of FDPs.

86

Increasing the number of pontics from three to four decreases the load-bearing capacity by nearly half. Increasing the number of pontics further would probably weaken the FDPs even more, but more research must be conducted to be able to predict the results.

The shape of a split-file designed all-ceramic reconstruction calls for a different dimension protocol compared to traditionally shaped ones as the split-file design leads to sharp proximal indentations acting as fractural impressions, thus decreasing the overall strength.

Framework design is a crucial factor for the load-bearing capacity of an all-ceramic FDP. The state-of-the-art design is preferable since split-file designed cores require a cross-sectional connector area that is at least 42% larger if they are to have the same load-bearing capacity as state-of-the-art designed cores.

Analysis of fracture patterns show differences between milled veneers and overpressed or built-up veneers; the milled ones show numerically more veneer cracks, and the other groups only show complete connector fractures.

87

ACKNOWLEDGEMENTS

I would like to express my deepest gratitude to all who have made this thesis possible, giving me all the inspiring, energy and generously support, with special thanks to:

Professor Ann Wennerberg, my supervisor; always have motivated and encouraging me to start and complete this work, sharing your expert guidance in academic experience, writing and thinking in the field of research, always with positive attitude, helpful and always being available.

Associated professor Per Vult von Steyern, my co-supervisor; have motivated me during the years and continue the journey, for sharing your academic thinking and knowledge with me and for encouraging and guiding me.

Ewa Linderoth my co-author and friend for sharing your fantastic knowledge, kindness and your profession in writing and proof-reading. Thanks for all long discussions, always have motivated and encouraging, without you this would not be possible.

Michael Braian my friend and researcher, I will never forget your first words at our very first introduction we meet. For always sharing your knowledge, pedagogy and engagement, your never-ending enthusiasm and giving me wise advice. For supporting me and always have motivated and encouraging me to start university and complete this work. For the helping to design the front layout of this thesis.

88

Arman Ameri for sharing your knowledge and making the tooth preparations.

Evaggelia Papia for sharing your knowledge, for all support and always making time for discussion.

Zdravko Bahat for publishing study I and always be friendly and for supporting.

Håkan Fransson; Watching the TC and always be around and checking up on me.

Margareta Molin Thoren; for valuable comments on the research program during my half-time seminar.

All the colleagues at the Faculty of Odontology, Malmö University and especially the entire staff of the Department of Prosthodontics and the Department of the Material Science and Technology; thanks for all help, support and care.

All the staff at Faculty Service; always be to hands and making work.

My fantastic friends; for all you love and support.

My wife and dear friend for your love, care and always being there for me and support, professional advice and being an excellent mother to our beautiful child.

My family, my mother, grandmother, brother and sisters for all your love, care and always being there for me and support.

This work has been made possible by generous support and grants from:The Faculty of Odontology, Malmö University Nobel Biocare in Sweden, special thanks Håkan Wiberg Ivoclar Vivadent in Sweden, special thanks to Pelle von Wowern and Christer HardingDenthouse AB in Sweden, special thanks to Tomas Lind Cosmodent AB in Sweden, special thanks to Przemek Seweryniak RH Dental ApS in Denmark

89

REFERENCES

1. WHO. International classification of functioning, disability and health (ICF). Ge-neva: WHO (2001). 2001.

2. Trulsson U, Engstrand P, Berggren U, Nannmark U, Branemark PI. Edentulousness and oral rehabilitation: experiences from the patients’ perspective. Eur J Oral Sci. 2002; 110: 417-424.

3. Bagewitz IC. Prosthodontics, care utilization and oral health-related quality of life. Swed Dent J Suppl. 2007; (185): 7-81.

4. Fardal O, Linden GJ. Long-term outcomes for cross-arch stabilizing bridges in periodontal maintenance patients-a retrospective study. J Clin Periodontol. 2010; 37: 299-304.

5. Bonding to oxide ceramics-laboratory testing versus clinical outcome. Dent Mater. 2015 Jan;31(1):8-14, Epub 2014 Jul 21.

6. Papia E, Larsson C, du Toit M, Vult von Steyern P. Bonding between oxide ceramics and adhesive cement systems: a systematic review. J Biomed Mater Res B Appl Biomater. 2014; 102: 395-413.

7. Milleding P. (ed) Traditional prosthodontic preparations. In: Nilner K, Karlsson S, Dahl B L (editors). A textbook of fixed prosthodontics: the scandinavian approach. 2 ed. Stockholm: Gothia fortbildning; 2013: 186-204.

8. Takuma Y, Nomoto S, Sato T, Sugihara N. Effect of framework design on fracture resistance in zirconia 4-unit all-ceramic fixed partial dentures. Bull Tokyo Dent Coll. 2013; 54: 149-156.

9. Raigrodski AJ, Yu A, Chiche GJ, Hochstedler JL, Mancl LA, Mohamed SE. Clinical efficacy of veneered zirconium dioxide-based posterior partial fixed dental prostheses: five-year results. J Prosthet Dent. 2012; 108: 214-222.

10. Kamposiora P, Papavasiliou G, Bayne SC, Felton DA. Stress concentration in all-ceramic posterior fixed partial dentures. Quintessence Int. 1996; 27: 701-706.

90

11. Oh WS, Anusavice KJ. Effect of connector design on the fracture resistance of all-ceramic fixed partial dentures. J Prosthet Dent. 2002; 87: 536-542.

12. Oilo M, Kvam K, Gjerdet NR. Simulation of clinical fractures for three different all-ceramic crowns. Eur J Oral Sci. 2014; 122: 245-250.

13. Aydin C, Yilmaz H, Bankoglu M. A single-tooth, two-piece zirconia implant located in the anterior maxilla: a clinical report. J Prosthet Dent. 2013; 109: 70-74.

14. Andreiotelli M, Wenz HJ, Kohal RJ. Are ceramic implants a viable alternative to titanium implants? A systematic literature review. Clin Oral Implants Res. 2009; 20 Suppl 4: 32-47.

15. Vult von Steyern P. Dental ceramics in clinical practice. Milleding P. Traditional prosthodontic preparations. In: Nilner K, Karlsson S, Dahl B L (ed). A textbook of fixed prosthodontics: the scandinavian approach. 2 ed. Stockholm: Gothia fortbildning; 2013: 205-222.

16. Anusavice KJ. Dental Ceramics. In: Anusavice KJ (ed). Phillipsʼ science of dental materials. St. Louis Saunders, 2013: 418-473.

17. Richerson DW. Time, temperature and enviromental effects on properties. In: Richerson DW (ed). Modern ceramic engineering: properties, processing, and use in design. Boca Raton, FL: CRC Taylor & Francis, 2006: 243-283.

18. Cesar PF, Yoshimura HN, Miranda WG,Jr, Miyazaki CL, Muta LM, Rodrigues Filho LE. Relationship between fracture toughness and flexural strength in dental porcelains. J Biomed Mater Res B Appl Biomater. 2006; 78: 265-273.

19. Scherrer SS, Kelly JR, Quinn GD, Xu K. Fracture toughness (KIc) of a dental porcelain determined by fractographic analysis. Dent Mater. 1999; 15: 342-348.

20. Cesar PF, Soki FN, Yoshimura HN, Gonzaga CC, Styopkin V. Influence of leucite content on slow crack growth of dental porcelains. Dent Mater. 2008; 24: 1114-1122.

21. Rosa V, Yoshimura HN, Pinto MM, Fredericci C, Cesar PF. Effect of ion exchange on strength and slow crack growth of a dental porcelain. Dent Mater. 2009; 25: 736-743.

22. Richerson DW. Design approches. In: Richerson DW (ed). Modern ceramic engineering: properties, processing, and use in design. Boca Raton, FL: CRC Taylor & Francis 2006: 581-594.

23. McLean JW. Evolution of dental ceramics in the twentieth century. J Prosthet Dent. 2001; 85: 61-66.

24. McLaren EA CP. Ceramics in dentistry-part I: classes of materials. Inside Dentistry. 2009; 5 (9): 94-105.

91

25. Kelly JR, Benetti P. Ceramic materials in dentistry: historical evolution and current practice. Aust Dent J. 2011; 56 Suppl 1: 84-96.

26. International Standards Organization. (ed) Dentistry - Ceramic materials Geneva, Switzerland: European Committee for Standardization; 2008. ISO 6872:2008 .

27. Blatz MB, Sadan A, Arch GH,Jr, Lang BR. In vitro evaluation of long-term bonding of Procera AllCeram alumina restorations with a modified resin luting agent. J Prosthet Dent. 2003; 89: 381-387.

28. Höland W, Schweiger M, Watzke R, Peschke A, Kappert H. Ceramics as biomaterials for dental restoration. Expert review of medical devices. 2008; 5: 729-45.

29. Reich S, Endres L, Weber C, Wiedhahn K, Neumann P, Schneider O et al. Three-unit CAD/CAM-generated lithium disilicate FDPs after a mean observation time of 46 months. Clin Oral Investig. 2014 Dec;18(9):2171-8.

30. Kern M, Sasse M, Wolfart S. Ten-year outcome of three-unit fixed dental prostheses made from monolithic lithium disilicate ceramic. J Am Dent Assoc. 2012; 143: 234-240.

31. Gorman CM, Horgan K, Dollard RP, Stanton KT. Effects of repeated processing on the strength and microstructure of a heat-pressed dental ceramic. J Prosthet Dent. 2014 Dec;112(6):1370-6.

32. Li RWK, Chow T, Matinlinna J. Ceramic dental biomaterials and CAD/CAM technology: State of the art. J Prosthodont Res. 2014 Oct;58(4): 208-16.

33. Denry I, Holloway JA. Ceramics for Dental Applications: A Review. Materials. 2010; 3: 351-368.

34. Kaiser M, Wasserman A, Strub JR. Long-term clinical results of VITA In-Ceram Classic: a systematic review. Schweiz Monatsschr Zahnmed. 2006; 116: 120-128.

35. Vult vS, Jonsson O, Nilner K. Five-year evaluation of posterior all-ceramic three-unit (In-Ceram) FPDs. Int J Prosthodont. 2001; 14: 379-384.

36. Zarone F, Russo S, Sorrentino R. From porcelain-fused-to-metal to zirconia: clinical and experimental considerations. Dent Mater. 2011; 27: 83-96.

37. Filser F, Kocher P, Weibel F, Luthy H, Scharer P, Gauckler LJ. Reliability and strength of all-ceramic dental restorations fabricated by direct ceramic machining (DCM). Int J Comput Dent. 2001; 4: 89-106.

38. McLaren EA, Terry DA. CAD/CAM systems, materials, and clinical guidelines for all-ceramic crowns and fixed partial dentures. Compend Contin Educ Dent. 2002; 23: 637-41, 644, 646 passim; quiz 654.

92

39. Lin WS, Metz MJ, Pollini A, Ntounis A, Morton D. Digital data acquisition for a CAD/CAM-fabricated titanium framework and zirconium oxide restorations for an implant-supported fixed complete dental prosthesis. J Prosthet Dent. 2014 Dec;112(6):1324-9.

40. Wimmer T, Hostettler J, Beuer F, Stawarczyk B. Load-bearing capacity of soldered and subsequently veneered 4-unit zirconia FDPs. J Mech Behav Biomed Mater. 2013; 23: 1-7.

41. Koutayas S, Vagkopoulou T, Pelekanos S, Koidis P, Strub J. Zirconia in dentistry: part 2. Evidence-based clinical breakthrough. The European journal of esthetic dentistry. 2009; 4: 348-80.

42. Johansson C, Kmet G, Rivera J, Larsson C, Vult Von Steyern P. Fracture strength of monolithic all-ceramic crowns made of high translucent yttrium oxide-stabilized zirconium dioxide compared to porcelain-veneered crowns and lithium disilicate crowns. Acta Odontol Scand. Epub 2013 Jul 18, 2014 Feb;72(2):145-53.

43. Zhang Y. Making yttria-stabilized tetragonal zirconia translucent. Dental materials. 2014; 30: 1195-203.

44. Piconi C, Maccauro G. Zirconia as a ceramic biomaterial. Biomaterials. 1999; 20: 1-25.

45. Guazzato M, Quach L, Albakry M, Swain MV. Influence of surface and heat treatments on the flexural strength of Y-TZP dental ceramic. J Dent. 2005; 33: 9-18.

46. Lin J, Shinya A, Gomi H, Shinya A. Finite element analysis to compare stress distribution of connector of lithia disilicate-reinforced glass-ceramic and zirconia-based fixed partial denture. Odontology. 2012; 100: 96-99.

47. Quinn GD, Studart AR, Hebert C, VerHoef JR, Arola D. Fatigue of zirconia and dental bridge geometry: Design implications. Dent Mater. 2010; 26: 1133-1136.

48. Sikalidis C. Application of Zirconia in Dentistry: Biological, Mechanical and Optical Considerations. Boco Raton, FL: Taylor & Francis Group, LLC;. 2011: 397-420.

49. Wang G, Zhang S, Bian C, Kong H. Fracture mechanics analyses of ceramic/veneer interface under mixed-mode loading. J Mech Behav Biomed Mater. 2014; 39: 119-128.

50. Luthardt RG, Holzhuter MS, Rudolph H, Herold V, Walter MH. CAD/CAM-machining effects on Y-TZP zirconia. Dent Mater. 2004; 20: 655-662.

51. Piconi C, Labanti M, Magnani G, Caporale M, Maccauro G, Magliocchetti G. Analysis of a failed alumina THR ball head. Biomaterials. 1999; 20: 1637-1646.

93

52. Richard H. J. Hannink, Patrick M. Kelly, Barry C. Muddle. Transformation Toughening in Zirconia-Containing Ceramics. Journal of the American Ceramic Society. March 2000; 83: 461-487.

53. Lughi V, Sergo V. Low temperature degradation -aging- of zirconia: A critical review of the relevant aspects in dentistry. Dental materials. 2010; 26: 807-20.

54. Chevalier J, Loh J, Gremillard L, Meille S, Adolfson E. Low-temperature degradation in zirconia with a porous surface. Acta Biomater. 2011; 7: 2986-2993.

55. Swain MV. Impact of oral fluids on dental ceramics: what is the clinical relevance? Dent Mater. 2014; 30: 33-42.

56. Gonzaga C, Cesar P, Miranda W, Yoshimura H. Slow crack growth and reliability of dental ceramics. Dental materials. 2011; 27: 394-406.

57. Chevalier J. What future for zirconia as a biomaterial? Biomaterials. 2006; 27: 535-43.

58. Fabbri P, Piconi C, Burresi E, Magnani G, Mazzanti F, Mingazzini C. Lifetime estimation of a zirconia-alumina composite for biomedical applications. Dent Mater. 2014; 30: 138-142.

59. Ebeid K, Wille S, Hamdy A, Salah T, El Etreby A, Kern M. Effect of changes in sintering parameters on monolithic translucent zirconia. Dental materials. 2014 2014 Dec;30(12):e419-24. doi: 10.1016/j.dental.2014.09.003. Epub 2014 Sep 26.

60. Jing Z, Ke Z, Yihong L, Zhijian S. Effect of Multistep Processing Technique on the Formation of Micro-defects and Residual Stresses in Zirconia Dental Restorations. J Prosthodont. 2014 Apr;23(3):206-12.

61. Kim M, Ahn J, Kim J, Kim H, Kim W. Effects of the sintering conditions of dental zirconia ceramics on the grain size and translucency. The journal of advanced prosthodontics. 2013; 5: 161-6.

62. Oh G, Yun K, Lee K, Lim H, Park S. Sintering behavior and mechanical properties of zirconia compacts fabricated by uniaxial press forming. The journal of advanced prosthodontics. 2010; 2: 81-7.

63. Richerson DW. Densification. In: Richerson DW (ed). Modern ceramic engineering: properties, processing, and use in design. Boca Raton, FL: CRC Taylor & Francis 2006: 477-527.

64. J. Muñoz-Saldaña, H. Balmori-Ramírez, D. Jaramillo-Vigueras, T. Iga and G. A. Schneider. Mechanical properties and low-temperature aging of tetragonal zirconia polycrystals processed by hot isostatic pressing. Journal of Materials Research. 2003; 18: 2415-2426.

65. Boitelle P, Mawussi B, Tapie L, Fromentin O. A systematic review of CAD/CAM fit restoration evaluations. J Oral Rehabil. 2014; 41: 853-74.

94

66. Della Bona A, Nogueira AD, Pecho OE. Optical properties of CAD-CAM ceramic systems. J Dent. 2014; 42: 1202-1209.

67. Denry I. How and when does fabrication damage adversely affect the clinical performance of ceramic restorations? Dent Mater. 2013; 29: 85-96.

68. Curtis A, Wright A, Fleming GJP. The influence of surface modification techniques on the performance of a Y-TZP dental ceramic. J Dent. 2006; 34: 195-206.

69. Miyazaki T, Hotta Y. CAD/CAM systems available for the fabrication of crown and bridge restorations. Aust Dent J. 2011; 56 Suppl 1: 97-106.

70. Dittmer MP, Kohorst P, Borchers L, Stiesch-Scholz M. Finite element analysis of a four-unit all-ceramic fixed partial denture. Acta Biomater. 2009; 5: 1349-1355.

71. Onodera K, Sato T, Nomoto S, Miho O, Yotsuya M. Effect of connector design on fracture resistance of zirconia all-ceramic fixed partial dentures. Bull Tokyo Dent Coll. 2011; 52: 61-67.

72. Esquivel Upshaw J, Clark A, Shuster J, Anusavice K. Randomized clinical trial of implant-supported ceramic-ceramic and metal-ceramic fixed dental prostheses: preliminary results. Journal of prosthodontics. 2014; 23: 73-82.

73. Plengsombut K, Brewer JD, Monaco EA,Jr, Davis EL. Effect of two connector designs on the fracture resistance of all-ceramic core materials for fixed dental prostheses. J Prosthet Dent. 2009; 101: 166-173.

74. Schmitter M, Mussotter K, Rammelsberg P, Gabbert O, Ohlmann B. Clinical performance of long-span zirconia frameworks for fixed dental prostheses: 5-year results. J Oral Rehabil. 2012; 39: 552-7.

75. Harshitha Gowda BH, Satish Babu CL. Connector design in a long-span-fixed dental prosthesis: a three-dimensional finite element analysis. Indian Journal of Dental Research. 2013; 24: 178-82.

76. Shillingburg HT, Jr., Sather DA, Wilson EL, Jr., Cain JR, Mitchell DL, Blanco LJ et al. All-Ceramic Restorations. 2012; 23: 425-446.

77. Koenig V, Vanheusden AJ, Le Goff SO, Mainjot AK. Clinical risk factors related to failures with zirconia-based restorations: an up to 9-year retrospective study. J Dent. 2013; 41: 1164-1174.

78. Sailer I, Feher A, Filser F, Gauckler LJ, Luthy H, Hammerle CH. Five-year clinical results of zirconia frameworks for posterior fixed partial dentures. Int J Prosthodont. 2007; 20: 383-388.

79. Larsson C, Holm L, Lovgren N, Kokubo Y, Vult vS. Fracture strength of four-unit Y-TZP FPD cores designed with varying connector diameter. An in-vitro study. J Oral Rehabil. 2007; 34: 702-709.

80. Sundh A, Sjogren G. Fracture resistance of all-ceramic zirconia bridges with differing phase stabilizers and quality of sintering. Dent Mater. 2006; 22: 778-784.

95

81. Rosentritt M, Behr M, Scharnagl P, Handel G, Kolbeck C. Influence of resilient support of abutment teeth on fracture resistance of all-ceramic fixed partial dentures: an in vitro study. Int J Prosthodont. 2011; 24: 465-8.

82. Mahmood D, Linderoth E, Vult von Steyern P, Wennerberg A. Fracture strength of all-ceramic (Y-TZP) three- and four-unit fixed dental prostheses with different connector design and production history. Swed Dent J. 2013; 37, 4: 179-187.

83. Raigrodski AJ, Hillstead MB, Meng GK, Chung KH. Survival and complications of zirconia-based fixed dental prostheses: a systematic review. J Prosthet Dent. 2012; 107: 170-177.

84. Webber B, McDonald A, Knowles J. An in vitro study of the compressive load at fracture of Procera AllCeram crowns with varying thickness of veneer porcelain. J Prosthet Dent. 2003; 89: 154-60.

85. Wang G, Zhang S, Bian C, Kong H. Effect of Thickness Ratio on Load-Bearing Capacity of Bilayered Dental Ceramics. Journal of prosthodontics. 2014 2014 Jun 11. doi: 10.1111/jopr.12169. [Epub ahead of print].

86. Agustín Panadero R, Román Rodríguez J, Ferreiroa A, Solá Ruíz M, Fons Font A. Zirconia in fixed prosthesis. A literature review. Journal of Clinical and Experimental Dentistry. 2014 Feb 1;6(1):e66-73. doi: 10.4317/jced.51304. eCollection 2014

87. Mollers K, Patzold W, Parkot D, Kirsten A, Guth JF, Edelhoff D et al. Influence of connector design and material composition and veneering on the stress distribution of all-ceramic fixed dental prostheses: a finite element study. Dent Mater. 2011 Aug;27(8):e171-5.

88. Fernandes CP, Glantz PO, Nilner K. On the accuracy of some in vitro models for mechanical studies of maxillary removable partial dentures. Dent Mater. 2003; 19: 127-136.

89. Fischer J, Stawarczyk B, Hammerle CH. Flexural strength of veneering ceramics for zirconia. J Dent. 2008; 36: 316-321.

90. Ovri JEO. A parametric study of the biaxial strength test for brittle materials. Mat Chem Phys. 2000: 661-665.

91. Denry I, Kelly JR. Emerging Ceramic-based Materials for Dentistry. J Dent Res. 2014 Dec;93(12):1235-42.

92. Quinn JB, Quinn GD. A practical and systematic review of Weibull statistics for reporting strengths of dental materials. Dent Mater. 2010; 26: 135-147.

93. Bahat Z, Mahmood DJ, Vult von Steyern P. Fracture strength of three-unit fixed partial denture cores (Y-TZP) with different connector dimension and design. Swed Dent J. 2009; 33: 149-159.

94. Zhang Y, Sailer I, Lawn B. Fatigue of dental ceramics. J Dent. 2013; 41: 1135-47.

96

95. Taskonak B, Yan J, Mecholsky JJ,Jr, Sertgoz A, Kocak A. Fractographic analyses of zirconia-based fixed partial dentures. Dent Mater. 2008; 24: 1077-1082.

96. Mahmood DJ, Linderoth EH, Vult Von Steyern P. The influence of support properties and complexity on fracture strength and fracture mode of all-ceramic fixed dental prostheses. Acta Odontol Scand. 2011 Jul;69(4):229-37.

97. Wimmer T, Erdelt KJ, Eichberger M, Roos M, Edelhoff D, Stawarczyk B. Influence of abutment model materials on the fracture loads of three-unit fixed dental prostheses. Dent Mater J. 2014; 33: 717-724.

98. Morresi A, D’Amario M, Capogreco M, Gatto R, Marzo G, D’Arcangelo C et al. Thermal cycling for restorative materials: does a standardized protocol exist in laboratory testing? A literature review. Journal of the mechanical behavior of biomedical materials. 2014; 29: 295-308.

99. Nicolaisen M, Bahrami G, Finlay S, Isidor F. Comparison of fatigue resistance and failure modes between metal-ceramic and all-ceramic crowns by cyclic loading in water. J Dent. 2014 Dec;42(12):1613-20.

100. Murase T, Nomoto S, Sato T, Shinya A, Koshihara T, Yasuda H. Effect of Connector Design on Fracture Resistance in All-ceramic Fixed Partial Dentures for Mandibular Incisor Region. Bull Tokyo Dent Coll. 2014; 55: 149-155.

101. Luthy H, Filser F, Loeffel O, Schumacher M, Gauckler LJ, Hammerle CH. Strength and reliability of four-unit all-ceramic posterior bridges. Dent Mater. 2005; 21: 930-937.

102. Sundh A, Molin M, Sjogren G. Fracture resistance of yttrium oxide partially-stabilized zirconia all-ceramic bridges after veneering and mechanical fatigue testing. Dent Mater. 2005; 21: 476-482.

103. de França, Danilo Gonzaga B, Morais, Maria Helena S T, das Neves F, Barbosa GAS. Influence of CAD/CAM on the fit accuracy of implant-supported zirconia and cobalt-chromium fixed dental prostheses. J Prosthet Dent. 2015; 113: 22-8.

104. Chitmongkolsuk S, Heydecke G, Stappert C, Strub JR. Fracture strength of all-ceramic lithium disilicate and porcelain-fused-to-metal bridges for molar replacement after dynamic loading. Eur J Prosthodont Restor Dent. 2002; 10: 15-22.

105. Kheradmandan S, Koutayas SO, Bernhard M, Strub JR. Fracture strength of four different types of anterior 3-unit bridges after thermo-mechanical fatigue in the dual-axis chewing simulator. J Oral Rehabil. 2001; 28: 361-369.

97

106. Tinschert J, Natt G, Mautsch W, Augthun M, Spiekermann H. Fracture resistance of lithium disilicate-, alumina-, and zirconia-based three-unit fixed partial dentures: a laboratory study. Int J Prosthodont. 2001; 14: 231-238.

107. Vult vS, Kokubo Y, Nilner K. Use of abutment-teeth vs. dental implants to support all-ceramic fixed partial dentures: an in-vitro study on fracture strength. Swed Dent J. 2005; 29: 53-60.

108. Rinke S, Gersdorff N, Lange K, Roediger M. Prospective evaluation of zirconia posterior fixed partial dentures: 7-year clinical results. Int J Prosthodont. 2013; 26: 164-171.

109. Sax C, Hämmerle CHF, Sailer I. 10-year clinical outcomes of fixed dental prostheses with zirconia frameworks. Int J Comput Dent. 2011; 14: 183-202.

110. Sailer I, Gottnerb J, Kanelb S, Hammerle CH. Randomized controlled clinical trial of zirconia-ceramic and metal-ceramic posterior fixed dental prostheses: a 3-year follow-up. Int J Prosthodont. 2009; 22: 553-560.

111. Molin MK, Karlsson SL. Five-year clinical prospective evaluation of zirconia-based Denzir 3-unit FPDs. Int J Prosthodont. 2008; 21: 223-227.

112. Ioannis K, Stephan J, Michael R, Klaus B. Prospective evaluation of zirconia based tooth- and implant- supported fixed dental prostheses: 3-Year Results. J Dent. 2014 Jan;43(1):87-93.

113. Le M, Papia E, Larsson C. The clinical success of tooth- and implant-supported zirconia-based fixed dental prostheses. A systematic review. J Oral Rehabil. 2015 2015 Jan 10. doi: 10.1111/joor.12272. [Epub ahead of print].

114. Kanat B, Comlekoglu EM, Dundar-Comlekoglu M, Hakan Sen B, Ozcan M, Ali Gungor M. Effect of Various Veneering Techniques on Mechanical Strength of Computer-Controlled Zirconia Framework Designs. J Prosthodont. 2014 Aug;23(6):445-55.

115. Wiedhahn K. The impression-free Cerec multilayer bridge with the CAD-on method. Int J Comput Dent. 2011; 14: 33-45.

116. Andreiotelli M, Kamposiora P, Papavasiliou G. Digital data management for CAD/CAM technology. An update of current systems. Eur J Prosthodont Restor Dent. 2013; 21: 9-15.

117. Miyazaki T, Hotta Y, Kunii J, Kuriyama S, Tamaki Y. A review of dental CAD/CAM: current status and future perspectives from 20 years of experience. Dent Mater J. 2009; 28: 44-56.

118. Pereira GKR, Amaral M, Simoneti R, Rocha GC, Cesar PF, Valandro LF. Effect of grinding with diamond-disc and -bur on the mechanical behavior of a Y-TZP ceramic. Journal of the mechanical behavior of biomedical materials. 2014; 37: 133-40.

98

119. Aboushelib MN, de Kler M, van der Zel JM, Feilzer AJ. Effect of veneering method on the fracture and bond strength of bilayered zirconia restorations. Int J Prosthodont. 2008; 21: 237-240.

120. Beuer F, Steff B, Naumann M, Sorensen JA. Load-bearing capacity of all-ceramic three-unit fixed partial dentures with different computer-aided design (CAD)/computer-aided manufacturing (CAM) fabricated framework materials. Eur J Oral Sci. 2008; 116: 381-386.

121. WJ O’Brien. Appendix A, Elastic Moduluod, E. 2008: 330-335.

122. Ibrahim I, Mohamed F, Lavernia E. Particulate reinforced metal matrix composites a review. J Mater Sci. 1991; 26: 1137-1156.

123. NORDBERGS TEKNISKA AB. Materialdata ( POM ) [Web Page]. [ 2015 01/13].

124. Gale MS, Darvell BW. Thermal cycling procedures for laboratory testing of dental restorations. J Dent. 1999; 27: 89-99.

125. Vatali A, Kontonasaki E, Kavouras P, Kantiranis N, Papadopoulou L, Paraskevopoulos KK et al. Effect of heat treatment and in vitro aging on the microstructure and mechanical properties of cold isostatic-pressed zirconia ceramics for dental restorations. Dent Mater. 2014 Oct;30(10):e272-82. doi: 10.1016/j.dental.2014.05.017. Epub 2014 Jun 17.

126. Kohorst P, Dittmer MP, Borchers L, Stiesch-Scholz M. Influence of cyclic fatigue in water on the load-bearing capacity of dental bridges made of zirconia. Acta Biomater. 2008; 4: 1440-1447.

127. Yi Y, Kelly JR. Failure responses of a dental porcelain having three surface treatments under three stressing conditions. Dental materials. 2011; 27: 1252-8.

128. Rosentritt M, Behr M, van der Zel JM, Feilzer A. Approach for valuating the influence of laboratory simulation. Dental materials. 2009; 25: 348-52.

129. Bhowmick S, Melendez-Martinez JJ, Hermann I, Zhang Y, Lawn BR. Role of indenter material and size in veneer failure of brittle layer structures. J Biomed Mater Res B Appl Biomater. 2007; 82: 253-259.

130. Kelly JR, Tesk JA, Sorensen JA. Failure of all-ceramic fixed partial dentures in vitro and in vivo: analysis and modeling. J Dent Res. 1995; 74: 1253-1258.

131. Ritter JE. Critique of test methods for lifetime predictions. Dent Mater. 1995; 11: 147-151.

132. Zeng K, Oden A, Rowcliffe D. Evaluation of mechanical properties of dental ceramic core materials in combination with porcelains. Int J Prosthodont. 1998; 11: 183-189.

133. Zeng K, Oden A, Rowcliffe D. Flexure tests on dental ceramics. Int J Prosthodont. 1996; 9: 434-439.

99

134. Fischer H, Marx R. Fracture toughness of dental ceramics: comparison of bending and indentation method. Dental materials. 2002; 18: 12-9.

135. Kelly JR. Clinically relevant approach to failure testing of all-ceramic restorations. J Prosthet Dent. 1999; 81: 652-661.

136. Vult von Steyern P, Ebbesson S, Holmgren J, Haag P, Nilner K. Fracture strength of two oxide ceramic crown systems after cyclic pre-loading and thermocycling. J Oral Rehabil. 2006; 33: 682-689.

137. Kelly JR. Perspectives on strength. Dent Mater. 1995; 11: 103-110.

138. Tartaglia G, Sidoti E, Sforza C. Seven-year prospective clinical study on zirconia-based single crowns and fixed dental prostheses. Clin Oral Investig. 2014 Oct 12. [Epub ahead of print].

139. Lassila V, Holmlund I, Koivumaa KK. Bite force and its correlations in different denture types. Acta Odontol Scand. 1985; 43: 127-132.

140. Koc D, Dogan A, Bek B. Bite force and influential factors on bite force measurements: a literature review. Eur J Dent. 2010; 4: 223-232.

141. Bonakdarchian M, Askari N, Askari M. Effect of face form on maximal molar bite force with natural dentition. Arch Oral Biol. 2009; 54: 201-4.

142. Kelly JR. Ceramics in restorative and prosthetic dentistry. Annu Rev Mater Sci. 1997; 27: 443-468.

143. Tripathi G, A A P, Rajwadha N, Chhaparia N, Sharma A, Anant M. Comparative evaluation of maximum bite force in dentulous and edentulous individuals with different facial forms. Journal of clinical and diagnostic research. 2014 Sep;8(9):ZC37-40. doi: 10.7860/JCDR/2014/8801.4837. Epub 2014 Sep 20.

144. Bakke M. Bite Force and occlusion. Seminars in Orthodontics. 2006; 12, Issue 2, June: 120-126.

145. Sorrentino R, De Simone G, Tete S, Russo S, Zarone F. Five-year prospective clinical study of posterior three-unit zirconia-based fixed dental prostheses. Clin Oral Investig. 2012; 16: 977-985.

146. Salido M, Martinez Rus F, del Rio F, Pradies G, Ozcan M, Suarez M. Prospective clinical study of zirconia-based posterior four-unit fixed dental prostheses: four-year follow-up. Int J Prosthodont. 2012; 25: 403-9.

147. Oh W, Gotzen N, Anusavice KJ. Influence of connector design on fracture probability of ceramic fixed-partial dentures. J Dent Res. 2002; 81: 623-627.

148. Molin MK, Onesti MP, Petersson TB, Derand TB. Three-dimensional finite element analyses of all-ceramic posterior fixed partial dentures with different designs. Int J Prosthodont. 2007; 20: 89-91.

149. Lee J, Choi S, Kim M, Kim H, Kim Y, Shin S. Effect of span length on the fit of zirconia framework fabricated using CAD/CAM system. The journal of advanced prosthodontics. 2013; 5: 118-25.

100

150. Beuer F, Edelhoff D, Gernet W, Sorensen JA. Three-year clinical prospective evaluation of zirconia-based posterior fixed dental prostheses (FDPs). Clin Oral Investig. 2009; 13: 445-451.

151. Ambre MJ, Aschan F, von Steyern PV. Fracture Strength of Yttria-Stabilized Zirconium-Dioxide (Y-TZP) Fixed Dental Prostheses (FDPs) with Different Abutment Core Thicknesses and Connector Dimensions. J Prosthodont. 2013 Jul;22(5):377-82.

152. Scherrer SS, Cattani Lorente M, Yoon S, Karvonen L, Pokrant S, Rothbrust F et al. Post-hot isostatic pressing: a healing treatment for process related defects and laboratory grinding damage of dental zirconia? Dental materials. 2013; 29: e180-90.

153. Hmaidouch R, Müller W, Lauer H, Weigl P. Surface roughness of zirconia for full-contour crowns after clinically simulated grinding and polishing. International journal of oral science. 2014 Dec;6(4):241-6.

154. Preis V, Letsch C, Handel G, Behr M, Schneider-Feyrer S, Rosentritt M. Influence of substructure design, veneer application technique, and firing regime on the in vitro performance of molar zirconia crowns. Dent Mater. 2013 Jul;29(7):e113-21

155. Karimipour Saryazdi M, Sadid Zadeh R, Givan D, Burgess J, Ramp L, Liu P. Influence of surface treatment of yttrium-stabilized tetragonal zirconium oxides and cement type on crown retention after artificial aging. J Prosthet Dent. 2014; 111: 395-403.

156. Ferrari M, Vichi A, Zarone F. Zirconia abutments and restorations: From laboratory to clinical investigations. Dental materials. 2015 Jan 6. pii: S0109-5641(14)00669-1. doi: 10.1016/j.dental.2014.11.015. [Epub ahead of print].

157. Kolgeci L, Mericske E, Worni A, Walker P, Katsoulis J, Mericske Stern R. Technical complications and failures of zirconia-based prostheses supported by implants followed up to 7 years: a case series. Int J Prosthodont. 2014; 27: 544-52.

158. Nossair S, Aboushelib M, Morsi T. Fracture and Fatigue Resistance of Cemented versus Fused CAD-on Veneers over Customized Zirconia Implant Abutments. Journal of prosthodontics. 2015 Jan 5. doi: 10.1111/jopr.12253. [Epub ahead of print].

159. Guess PC, Zhang Y, Kim J, Rekow ED, Thompson VP. Damage and reliability of Y-TZP after cementation surface treatment. J Dent Res. 2010; 89: 592-6.

160. Christensen R, Ploeger B. A clinical comparison of zirconia, metal and alumina fixed-prosthesis frameworks veneered with layered or pressed ceramic: a three-year report. J Am Dent Assoc. 2010; 141: 1317-29.

101

161. Costa AKF, Borges ALS, Fleming GJP, Addison O. The strength of sintered and adhesively bonded zirconia/veneer-ceramic bilayers. J Dent. 2014; 42: 1269-76.

162. Flury S, Peutzfeldt A, Lussi A. Influence of surface roughness on mechanical properties of two computer-aided design/computer-aided manufacturing (CAD/CAM) ceramic materials. Oper Dent. 2012; 37: 617-624.

163. Schmitter M, Schweiger M, Mueller D, Rues S. Effect on in vitro fracture resistance of the technique used to attach lithium disilicate ceramic veneer to zirconia frameworks. Dent Mater. 2014; 30: 122-130.

164. Beuer F, Schweiger J, Eichberger M, Kappert HF, Gernet W, Edelhoff D. High-strength CAD/CAM-fabricated veneering material sintered to zirconia copings-a new fabrication mode for all-ceramic restorations. Dent Mater. 2009; 25: 121-128.

165. Marrelli M, Maletta C, Inchingolo F, Alfano M, Tatullo M. Three-point bending tests of zirconia core/veneer ceramics for dental restorations. Int J Dent. 2013;2013:831976. doi: 10.1155/2013/831976. Epub 2013 Feb 25.

166. Stawarczyk B, Ozcan M, Hammerle CH, Roos M. The fracture load and failure types of veneered anterior zirconia crowns: an analysis of normal and Weibull distribution of complete and censored data. Dent Mater. 2012; 28: 478-487.

167. Marchack BW, Sato S, Marchack CB, White SN. Complete and partial contour zirconia designs for crowns and fixed dental prostheses: a clinical report. J Prosthet Dent. 2011; 106: 145-152.

168. Daou E. The zirconia ceramic: strengths and weaknesses. Open Dentistry Journal. 2014; 8: 33-42.

169. Fischer J, Stawarczyk B, Tomic M, Strub JR, Hammerle CH. Effect of thermal misfit between different veneering ceramics and zirconia frameworks on in vitro fracture load of single crowns. Dent Mater J. 2007; 26: 766-772.

170. Swain MV. Unstable cracking (chipping) of veneering porcelain on all-ceramic dental crowns and fixed partial dentures. Acta Biomaterialia. 2009; 5: 1668-77.

171. Gostemeyer G, Jendras M, Dittmer MP, Bach FW, Stiesch M, Kohorst P. Influence of cooling rate on zirconia/veneer interfacial adhesion. Acta Biomater. 2010; 6: 4532-4538.

172. Meira JBC, Reis B, Tanaka C, Ballester R, Cesar P, Versluis A et al. Residual stresses in Y-TZP crowns due to changes in the thermal contraction coefficient of veneers. Dental materials. 2013; 29: 594-601.

173. Benetti P, Kelly JR, Sanchez M, Della Bona A. Influence of thermal gradients on stress state of veneered restorations. Dental materials. 2014; 30: 554-63.

102

174. Rues S, Kroger E, Muller D, Schmitter M. Effect of firing protocols on cohesive failure of all-ceramic crowns. J Dent. 2010; 38: 987-994.

175. Paula V, Lorenzoni F, Bonfante E, Silva, Nelson R F A, Thompson V, Bonfante G. Slow cooling protocol improves fatigue life of zirconia crowns. Dental materials. 2015 Feb;31(2):77-87.

176. Al Amleh B, Neil Waddell J, Lyons K, Swain M. Influence of veneering porcelain thickness and cooling rate on residual stresses in zirconia molar crowns. Dental materials. 2014; 30: 271-80.

177. Al-Amleh B, Lyons K, Swain M. Clinical trials in zirconia: a systematic review. J Oral Rehabil. 2010; 37: 641-652.

178. Ahlers MO, Mörig G, Blunck U, Hajtó J, Pröbster L, Frankenberger R. Guidelines for the preparation of CAD/CAM ceramic inlays and partial crowns. Int J Comput Dent. 2009; 12: 309-25.

179. Guess P, Schultheis S, Wolkewitz M, Zhang Y, Strub J. Influence of preparation design and ceramic thicknesses on fracture resistance and failure modes of premolar partial coverage restorations. J Prosthet Dent. 2013; 110: 264-73.

180. Benetti P, Pelogia F, Valandro L, Bottino M, Della Bona A. The effect of porcelain thickness and surface liner application on the fracture behavior of a ceramic system. Dental materials. 2011; 27: 948-53.

181. Rekow ED, Silva NR, Coelho PG, Zhang Y, Guess P, Thompson VP. Performance of dental ceramics: challenges for improvements. J Dent Res. 2011; 90: 937-952.

deyar jallal hadi mahmood - muep mahmood2... · deyar jallal hadi mahmood on core and bi-layered...

Documents