Australian Dental Journal
The official journal of the Australian Dental Association
Australian Dental Journal 2012; 57: 23–30
ADRF RESEARCH REPORT
doi: 10.1111/j.1834-7819.2011.01638.x
The all-ceramic, inlay supported fixed partial denture. Part 3.
Experimental approach for validating the finite element
analysis
MC Thompson,* CJ Field, MV Swain*
*Biomaterials, Faculty of Dentistry, The University of Sydney, New South Wales.
Aeromechanical Engineering, Faculty of Engineering, The University of Sydney, New South Wales.
ABSTRACT
In a previous study, the authors used a finite element analysis (FEA) to evaluate the stresses developed during the loading of
an all-ceramic, inlay supported fixed partial denture and compared it with the more traditional full crown supported
prosthesis. To date there has been little research into correlating the responses of the numerical model against physical
mechanical tests; such validation analysis is crucial if the results from the FEA are to be confidently relied upon. This study
reports on the experimental methods used to compare with the FEA and thereby to validate the predictive fracture behaviour
of the numerical model. This study also outlines the methods for manufacture and testing of the ceramic structure along
with observations of the fracture tests. In addition the procedure used for developing the FEA model for the test system is
outlined.
Keywords: All-ceramic bridge, inlay fixed partial denture, experimental design, finite element analysis.
Abbreviations and acronyms: CAD ⁄ CAM = computer-aided design ⁄ computer-aided milling; CR = composite resin; FEA = finite element
analysis; FPD = fixed partial denture; PDL = periodontal ligament; Y-TZP = yttria stabilized tetragonal zirconia.
(Accepted for publication 1 August 2011.)
INTRODUCTION
A comprehensive literature review into the ideal tooth
preparation design for a ceramic inlay has previously
been conducted by the authors;1 critical aspects such as
cavity depth, isthmus width and total occlusal convergence were analysed and optimized cavity geometry
formulated. A subsequent paper utilized the optimized
inlay preparation design on abutments supporting an
all-ceramic, three-unit fixed partial denture (FPD) with
the lower first molar as the pontic. The prosthesis form
design was improved to better distribute tensile forces
particularly in the gingival embrasure area which
displays peak tensile forces; this highly developed inlay
supported prosthesis was compared against the conventional full crown supported FPD via a finite element
analysis (FEA).2
The results from this numerical study predicted that
peak stresses in the inlay bridge are around 20% higher
than in the full crown supported bridge with von Mises
stresses peaking at about 730 MPa (the theoretical
ª 2012 Australian Dental Association
maximum strength of Y-ZTP is 900 to 1200 MPa; in
practice this is somewhat lower) when subjected to
theoretical average maximum bite force in the molar
region of 700 N. Maximum von Mises stresses
occurred at the loading sites; the embrasure areas and
the axio-pulpal line angle with maximum principal
stresses of 693 MPa being confined to the gingival
aspect of the embrasures. With careful and proper
design, it was concluded that the use of inlays as
support for an all-ceramic FPD could be successful.
Validation of the FEA is crucial if confidence in the
results is to be accepted. Development and manufacture of the physical prototype, the preparation and
mounting for testing and the testing procedure are all
crucial if the validation is to be accurate and truthful
in determining the genuineness of the FEA. There
appears to be little work comparing the FEA of
ceramic FPDs against robust experiments. Earlier
validation studies have been quite rudimentary in their
methodology, with little regard given to matching
elastic moduli or realistic geometry; later studies are
23
MC Thompson et al.
increasingly adopting realistic experimental designs
with greater consideration given to matching material properties. However, there appears to still be
absent accurate geometry, scale materials and loading
conditions.3–9
Realistic experiments are required in order to confirm
the validity of the FEA. Experiments relying upon
simple beam loading designs and flexural tests as an
approximation of the complex 3-D pattern of stresses
and strains acting upon ceramics in the FEA must be
considered highly misleading and not reflective of the
overall simulation6 or clinical situation. Additionally,
too little regard has been given to the need for matching
the modulus of elasticity of all components within the
experimental boundaries as seen in models utilizing
metal abutments;10–12 this is especially relevant to the
elastic modulus of the supporting structures which are
primarily responsible in the distribution of occlusal
stresses.1,13–16 Results from experiments failing to
adequately deal with clinical reality must therefore be
guarded in their applicability and not be interpreted as
being absolute in their validity.
Unlike metallic alloys and other elastic–inelastic
deforming solids, the strength of ceramics structures is
dependent upon the testing method, specimen size and
geometry, state of the surface, internal stresses and
preparation method; any and all of which can significantly influence the results.13 This is particularly true
where notching occurs (as in the gingival embrasure
area of a FPD) due to development of a complex stress–
strain state resulting from the constraint of the notch.14
Hence, if the data are to be reliable and validation true,
the simulation must be realistic and as accurate a
replication to the FEA as possible.
A particularly important issue concerns the close
matching of the elastic moduli of the systems being
commensurate with that of the relative strengths of the
systems being tested and outcomes desired. The anticipated strength of the fully sintered yttria stabilized
tetragonal zirconia (Y-TZP) is so high (700–1200 MPa)
that if utilized would either necessitate the use of even
stronger abutment material, thus introducing a fundamental mismatch of the critical elastic moduli of the
abutments, or maintaining the use of CR abutments
and risk testing the strength and bonding of the
teeth ⁄ acrylic base as opposed to the bridge material.
Choosing to use partially sintered zirconia is an
important and novel distinguishing feature of the
experimental methodology in this paper. Its use allows
the precise and easy replication of numerous prototypes, each behaving in an elastic brittle manner
identical to the fully sintered material whilst having a
flexural strength low enough to enable three-point
loading to test the stresses and strains upon the bridge
design and not the adhesion and deformation of the
supports. The physical properties of the partially
24
sintered material are available neither from the manufacturer nor in the literature so the process of measuring
flexural strength, toughness and elastic modulus has
been described.
The aim of this study was to outline the methods used
to produce and test exact bridge prototypes derived
from the FEA virtual model, the testing of the ceramic
structure, the determination of the mechanical properties of the material which hitherto has not been
examined in detail and observations of the fracture
tests. In addition, the procedure used for developing the
FEA model for the test system is outlined.
MATERIALS AND METHODS
Numerical model construction – the 3-D FEA
The construction of the FEA has been detailed in a
previous study.2 Table 1 details the material properties
used in the FEA, in particular the Young’s modulus of
the partially sintered zirconia which was determined via
the use of the UMIS 2000 nano-indentation instrument
developed at the CSIRO.
Experimental model construction – the prototype
Model teeth made from Filtek P60 composite resin
(CR) due to its class leading compressive strength
(395 MPa) and similarity of Young’s modulus to
dentine (12.5 GPa)18 were fabricated by use of a
sectioned silicone impression of a typical human lower
right second molar and second premolar which closely
matched the FEA abutment teeth. The mould was
incrementally packed with P60, producing anatomical
crown and root segments which were then bonded to
form intact teeth (Fig. 1).
The inlay supported bridge STL file from the FEA
was sent to a specialized ceramics laboratory which was
able to mill via a computer-aided design ⁄ computeraided milled (CAD ⁄ CAM) system dimensionally and
geometrically precise bridge replicas in ‘green’ or
partially sintered Y-TZP (Fig. 2a–2c). Partially sintered
zirconia was preferred over the fully sintered material
due to its lower strength, thus allowing the supporting
replica teeth to be fabricated from composite resin as
Table 1. Material properties required within the FEA
models.17 Young’s modulus for ceramic determined by
indentation testing
Material
Enamel
Dentine
PDL
FPD ceramic
Young’s
modulus (MPa)
Poisson’s
Ratio
84 100
18 600
70.3
15 650
0.20
0.31
0.45
0.28
Partially sintered Y-TZP
ª 2012 Australian Dental Association
Experimental approach to validate FEA
Fig 1. Completed CR teeth to be used for inlay abutments.
(a)
(b)
(c)
Greatest accuracy in comparing the FEA to the
prototype can only be achieved by utilizing a material
with an elastic modulus very similar to dentine.
The ability of the milling process to produce a precise
replica of the inlay bridge developed in the FEM was
demonstrated by superimposing a high contrast image
of the inlay over the computer model.
Oversized inlay preparations were made in the CR
abutments ready to accept their respective restorations.
This resulted in a gap of approximately 1 mm around
the entire periphery of the inlays.
A self-curing acrylic (PMMA) base was constructed
prior with the alveolus of the abutment teeth moulded
in place by repeated insertion and removal of the
individual roots until the acrylic was almost set; this
resulted in two half bases, one for each abutment. After
the acrylic base had fully cured, a layer of poly vinyl
siloxane (3M Imprint II) simulating the natural resilience of the periodontal ligament (PDL) was coated
onto the roots and the teeth seated individually into
their respective sockets. Excess poly vinyl siloxane was
removed.
Oversized inlay preparations were treated with phosphoric acid. The cavity preparation and inlay component of the bridge was treated with an adhesive (3M
Single Bond 2) which was applied so as to saturate the
porous ceramic structure, leaving a bondable surface
film. A bed of CR was placed in the preparations and the
bridge placed under pressure so as to allow ‘squeeze out’
of material and subsequent re-adaptation of all margins.
This was fully cured on the ramp cycle of an Optilux
501 curing light to minimize stresses.
Three inlay-tooth models were tested to failure on a
Shimadzu AG-50 kNE with incremental and peak loads
being recorded together with displacement. The CR
teeth and base survived all testing and was reused, thus
guaranteeing identical abutments and testing conditions. A 5 mm hardened stainless steel ball was placed
(Fig. 3) so as to achieve stable three-point contact in the
central fossa of the pontic. Loading was at a rate of
0.5 m per minute and continued until fracture.
Ceramic specimens testing
Fig 2. (a) Buccal view of inlay supported FPD. (b) Inferior view of
FPD. (c) Occlusal view of FPD.
opposed to a metal die. The load-to-failure testing of a
fully sintered Y-TZP bridge on composite resin teeth
would likely lead to their failure long before the bridge.
Additionally, the use of metal dies or replica teeth
would drastically alter the Young’s modulus of the
primary supporting structure, resulting in errors.
ª 2012 Australian Dental Association
Strength tests are singularly the most important source
of information regarding the performance of a ceramic
during loading. The tests provide a source of data
regarding how a material may perform in service with
data compiled through simple loading conditions with
careful measurements taken of stress and stain.
Mechanical testing of ceramic specimens was essential
as no data are available in the literature; this was in order
to determine the physical properties for the partially
sintered zirconia and in particular it’s flexural or tensile
strength and the Young’s modulus. The three-point
loading technique has been suggested as the preferable
25
MC Thompson et al.
Fig 3. The inlay-tooth complex ready for testing.
technique for tensile strength testing of brittle materials
such as ceramics.19–21
Two series of tests were conducted on different sized
specimens to determine fracture or flexural strength; the
first being 5 · 5 mm on 20 mm outer supports and
the second 2.5 · 2.5 mm on 10 mm outer supports.
The purpose of this was to account for the effect of
specimen size or volume upon the strength. Failure
statistics as predicted by the Weibull approach
describes the strength distribution of brittle materials
based on the ‘weakest link’ theory; failure at any flaw
leads to total failure of each individual sample. The
largest flaw present in each specimen becomes the
weakest link and predicates its survivability; therefore
increasing specimen size also statistically increases the
chance that a large flaw is present and thus larger
specimens are generally weaker. A further series of tests
were conducted to determine the fracture toughness of
the partially sintered zirconia by way of the SEVNB test
(single-edge V notch beam test) (Fig. 4).
Preparation of all the ceramic samples was by
diamond separating disc; samples were then carefully
lapped on 1000 grit silicone carbide paper, emphasizing
a longitudinal scratch pattern, and attention paid to
ensure all corners were rounded and free of any
chipping which may act as crack initiation sites.
Samples were measured with a Mitutoyo Micrometer
prior to testing, and crack length measured after
fracture in the case of the notched sample tests.
Determining the best way of obtaining notched
samples has always been slightly problematic for
material scientists; as notch width decreases so too
does the fracture toughness value KIC. Values have been
found to level off at when the crack tip radius is below
10 lm,22 but how to obtain such a fine and consistent
notch has always been difficult. A razor blade was
inserted into the prepared V notch and moved back and
forth to obtain an immeasurably fine notch radius.
Fracture strength is defined as the maximum tensile
stress in the surface of a specimen fractured in bending
(it must be remembered that stresses acting across the
height of a specimen will vary from maximum tension
on one side to maximum compression on the other).
For a rectangular cross section specimen on the threepoint bend apparatus, tensile or yield strength is:
ru;b ¼
Fl
4Wb
where ru,b denotes the maximum strength in bending, F
is the fracture load and l is the width of the outer
supports.24 Wb is the sectional modulus of the specimen and is given by the equation:
Wb ¼ dh2 =6
where d and h the width and height of the rectangular
specimen respectively. Pre-cracked specimens were
used on the same three-point bend apparatus with
samples taken from the same milled lot as the 5 · 5 mm
specimens. The critical stress intensity factor is given by
the equation:
KIC ¼
3Fl
Ya1=2
2h2 d
for a ⁄ h = q < 0.6 is
Y ¼ 1:93 3:07q þ 14:53q2 25:11q3
þ 25:80q4
for l=h ¼ 4
Y ¼ 1:96 2:75q þ 13:66q2 23:98q3
þ 25:22q4
Fig 4. Single-edge V notch beam test.
26
for l=h ¼ 8
where KIC is the fracture toughness of the material, F is
the breaking load, l is the width of the outer supports, h
and d the height and width of the rectangular specimen
respectively, a the length of the crack ⁄ notch and Y the
dependent variable factor of the crack.23 It should be
ª 2012 Australian Dental Association
Experimental approach to validate FEA
noted that the height of the samples relative to their
support viz. aspect ratio is considered important; we
kept all sample testing within the 1:4 ratio for standardization. Elastic modulus testing for brittle materials has historically been inaccurate via the traditional
static or bench-top testing methods,24,25 largely due to
the inability to measure the small degrees of strain
involved. Traditional bend testing would underestimate
the correct value due to the inherent degree of
compliance present in all large scale equipment, masking the microdimensions of strain involved (£200 lm
for the 2.5 mm2 sample). However, current quasi-static
testing techniques such as nanoindentation are able to
easily and accurately determine such properties.
The Ultra Micro Indentation System, UMIS-2000
(CSIRO, Campbell, Australia) is a nanoindentation
instrument for investigation of the properties of
the near-surface region of materials. Nanoindentation
is a procedure which is non-destructive, relatively quick
to perform and uniquely capable of characterizing the
physical properties of both thin films and small volume
samples. Being easy and the penetration accuracy of the
indenter being just a few nanometres, it is considered
the ideal method for determining not only material
hardness, but elastic modulus and may be used for
the measure of strain hardening, creep, fracture and
other properties of surface and near surface regions
with unique accuracy following submicron depths of
penetration.
The UMIS-2000 utilizes several indenters including
a spherical diamond, the Vickers square pyramid
indenter and the Berkovich triangular pyramidal
diamond indenter. Principally, the Berkovich indenter
is used for brittle materials with a tip radius typically
50–100 nm, increasing to 200 nm with use and face
angle 65.03, its shaper geometry and longer diagonals
more readily lends its self to finer measurement where
penetration depth is shallow.24
elastic modulus and Poisson’s ratio of the ceramic
samples was via nanoindentation testing (Table 1).
Millability and verification of the inlay contours
Computer-aided milling of the zirconia bridge from the
STL file was verified dimensionally by comparing
measurements taken from the FEM against the prototype; to ascertain and demonstrate the exactness of
the geometry produced, a high contrast photo of the
prototype (Fig. 5a) was superimposed on the FEM
image of the bridge (Fig. 5b) with the ‘merge’ function
of Photoshop (Adobe CS2 version 9). The result is
demonstrated by the excellent confluence in the outlines
of the two images, particularly in the embrasure areas
and axio-pulpal line angles.
Experimental model fracture patterns
Visual inspection of all three failed three-unit bridge
models show that there was at least a single fracture
developing from the gingival aspect of the weakest
connector, travelling obliquely through the body of the
pontic (Fig. 6). In two of the models, a second fracture
developed vertically through the opposing connector
resulting in the complete fragmentation and separation
of the pontic. The loading site had evidence of slight
(a)
(b)
RESULTS
Ceramic specimen mechanical properties
Three-point testing of the 5 · 5 mm samples (n = 8)
averaged 46.6 MPa of fracture load with a standard
deviation of nearly 6 MPa; the 2.5 · 2.5 mm samples
(n = 14) averaged 50 MPa with a standard deviation of
just over 1 MPa. Testing of the SEVNB samples (n = 8)
demonstrated excellent consistency in the results with
an average fracture toughness value of 0.456 MPam
and standard deviation of 0.076 MPam. Note well
that the strain values given must not be regarded too
seriously as these micrometre dimensions are well
within the ability of the Shimadzu to resolve accurately
and hence unable to provide a precise Young’s modulus. The preferred option for the determination of the
ª 2012 Australian Dental Association
(c)
Fig 5. (a) High contrast photo of the buccal aspect of the zirconia
inlay. (b) FEM image of the buccal aspect of the zirconia inlay.
(c) Superimposition of Fig. 5a and 5b displaying the excellent
confluence of their outlines.
27
MC Thompson et al.
Fig 6. Three experimental models displaying fracture patterns.
fragmentation from where the ball contacted the fine
edge of the oblique fracture as a result of the
compressive nature of the damage.
DISCUSSION
The use of partially sintered zirconia milled directly
from the FEA STL file is a unique experimental
procedure. Utilizing a ceramic much weaker than the
supporting material has allowed the use of CR for the
abutments which has the advantage of not only
possessing a Young’s modulus close to that of dentine,
but the reuse of the original models, thus maintaining
consistency across all tests. Problems associated with
the production of numerous identical prototypes or
replicas are also overcome with milling directly from
the STL file as opposed to the more common method of
producing the FEM after the experimental model has
been made. As no mechanical data are available, threepoint bend tests were necessary in order to determine
the fracture strength of the partially sintered zirconia.
The fracture strength range was determined to be
46.6 to 50 MPa which is low enough to not risk
28
damage to the CR models during testing, whilst
sufficiently high to allow the routine handling and
seating of the bridge and most importantly accurate
milling without fear of chipping or crumbling of the
finer details. Milling was performed with a 0.5 mm
tungsten carbide cutter to an accuracy of within 10 lm,
leaving superb surface detail with no visible rippling or
tooling marks in the critical embrasure areas. As an
exercise to satisfy any doubts regarding the exactness of
the overall geometric outline of the milled bridge
compared to the numerical simulation, the superimposition of the two images demonstrates superb confluence in their overall shapes (Fig. 5c). Any perceivable
differences can be accounted for by the parallax error
and distortion from the cameras wide angle lens in
macro mode.
The lower figure of 46.6 MPa for fracture strength
tests on the three-point test with 20 mm outer support
compared to 50 MPa with the 10 mm outer supports is
as anticipated and in accordance with the Weibull
prediction of larger samples being somewhat weaker
than smaller volume samples.
Measurement of the strength of ceramics is known to
produce significant scatter in the results, this phenomenon is due to the random positioning of flaws present
throughout the structure, varying in size, location and
orientation. This necessitates a statistical approach to
describing quantitatively the data produced; hence
Weibull analysis.
Qualitatively, the site of crack initiation, direction,
pattern of growth are in excellent agreement with that
produced by other authors via bench testing26 and
state-of-the-art numerical crack simulation techniques.27 However, it must be emphasized that the
testing in this study, as in the authors’ previous paper,
differed crucially in that the anchorage of the pontic
was via an inlay as opposed to full crown. Research
articles concerning inlay supported fixed bridges are
very few, especially where the material is all-ceramic.
Of significance is where authors have sought to study
the fracture behaviour of ceramic bridges (all crown
supported); the observable results have been very
similar though the support of the pontic differs (inlay
vs. crown). It may be concluded that if the bond is
excellent and does not fail, then the fracture initiation
and progression as predicted by the FEA is the same in
both cases and not through the weak isthmus region of
the inlay as is common with Class II restorations.1
A recent two-dimensional (2-D) FEA analysis on the
stress distribution in inlay supported bridges28 concludes that regardless of material used for the inlay
bridge, the stress patterns are remarkably similar with
tensile peaks at the gingival aspect of the connectors,
and a compressive zone within the body of the pontic.
However, the 2-D FEA in the study did not display the
high compressive forces evident in our current 3-D FEA
ª 2012 Australian Dental Association
Experimental approach to validate FEA
at the occlusal aspect of the connectors (as would be
expected from the Law of Beams) and the loading of the
ball indenter, nor was it able to show the subtle
transition in stresses within the body of the inlay and
pontic. Fracture behaviour from the 2-D numerical
analysis could be predicted as being through the
connector or through the inlay; very little additional
information is given to confidently make any other
prediction as a result of the 2-D FEA nor were the
results validated against bench-top tests.
Fracture paths as predicted by the locations of the
highest maximum principal stresses from the FEA
model in this study can be clearly seen in Fig. 7a and
from the von Mises in Fig. 7b. This coincides with a
‘zone of failure’ extending between the peak tensile
stresses developing at the mesial and ⁄ or distal gingival
embrasures, and then extending to the occlusal loading
site. Failure under Model 1 (Fig. 6) depicts the crack
front propagating from the more intensely stressed of
the two gingival embrasures, extending to the occlusal
loading site. Contact damage on the occlusal of the
pontic developed as the crack tip approached the
loading site and encountered the Hertzian stresses
resulting from the loading ball. Models 2 and 3
(Fig. 6) show the fracture path developing from both
connector sites, resulting in an obliquely travelling
fracture path through the pontic, together with a
vertical crack extending through the opposite inlaypontic connector. Contact damage was also evident on
the occlusal surface of the pontic.
The initial crack development on the experimental
models occurred at the gingival aspect of the weakest
connector when the critical stress level is reached; this
accurately reflects the peak tensile stresses displayed by
(a)
(b)
Fig 7. (a) Predicted zones of failure from the maximum principle
stresses. (b) Predicted zones of failure from the peak von Mises
contours.
ª 2012 Australian Dental Association
the FEA. The formation and progression of cracks
within the ceramic takes place so as to maximize the rate
of energy release and minimize the strain within the
system; hence the direction of cracking necessarily
follows a path towards the highly stressed loading site
on the occlusal surface of the pontic. The development
of new cracks at the central fossa of the pontic originates
from the blunt contact damage of the loading ball which
is also anticipated from the FEA by very high von Mises
stresses on the pontic’s occlusal (see Fig. 1 in Thompson
et al.2). Initiation of the second fracture of the adjacent
connector as depicted in Cases 2 and 3 reflects the
equally high peak stresses developing at this gingival
surface. Cracking possibly follows closely after the first
connector has partially fractured and due to the inferior
rotation of the major pontic segment, leads to the
development of a vertical rather than another oblique
fracture. This process of crack initiation and progression is in an effort to release and redress the imbalance
of energy in the stress field as suggested by Irwin.23
As a comparison of the current FEA of the partially
sintered zirconia with the FEA the authors’ previous
paper,2 it is interesting to note that with full sintering
comes a 12 times increase in the stiffness of the material
as measured by the Young’s modulus. Even so, the
stress contours are remarkably similar, differing only
slightly on the quantum of the stresses resulting from
the 200 N load (209.9 MPa vs. 190 MPa for the von
Mises stress and 198 MPa vs. 191 MPa for the
maximum principle stress). Primarily, the largest discrepancy is observed on the degree to which the stresses
permeate or are transferred to the supporting inlays,
with critically high stress peaks occurring at the axiopulpal line angle of the fully sintered bridges but only
moderate stresses with the partially sintered bridge
reflecting the materials increased flexibility.
An encouraging recent study validated the results of
mechanical tests of milled ceramic bars against an
FEM, a process reversed from this current study.5 This
study demonstrates that under simplistic conditions, the
results of FEAs cannot only be precise but indeed more
reliable. In comparison, our current study was conducted on accurate models of genuine prosthesis and
their supporting structures, significantly more complex
in geometry and behaviour than simple ceramic bars.
Thus, the resultant stresses developed in the system is
affected by a wide range of variables, some of which are
difficult to quantify at present.
Our next study will correlate the predicted strength
values from the FEA against detailed stress and
statistical analysis of the inlay ceramic bridge and
ceramic samples in what is essentially the quantitative
validation. To date we have not been able to find any
stress figures derived from FEA of dental bridges that
have been successfully validated against experimental
models. There exist validation of simple bars with
29
MC Thompson et al.
numerical models but none relating to the complexity
of geometry existing with real FPDs.
8. Oh WS, Park JM, Anusavice KJ. Fracture toughness of a hotpressed core ceramic based on fractographic analysis of fractured
ceramic FPDs. Int J Prosthodont 2003;16:135–140.
CONCLUSIONS
9. Pietrabissa R, Contro R, Quaglini V, Soncini M, Gionso L,
Simion M. Experimental and computational approach for the
evaluation of the biomechanical effects of dental bridge misfit.
J Biomech 2000;33:1489–1495.
This study outlines the methodology used to produce
and test the bridge prototypes and qualitatively validates
the results of the FEA detailed in an earlier paper.2 The
basis for using partially sintered zirconia as a model
material is justified. The mechanical properties including
flexural strength and fracture toughness of the partially
sintered zirconia blocks has been measured. It is shown
that using resin based composite as a model for the
abutment teeth and partially sintered zirconia ceramic
for the bridge enabled us to evaluate the fracture
response of ball loading such a structure. The results
demonstrate excellent congruence between the fracture
behaviour predicted by the peak principal stresses from
the FEA. Initiation sites and propagation path of the
cracks established in the physical models are in agreement with the results obtained from the FEA. The
preliminary results reported will be extended to explore
additional aspects of the role of the boundary conditions,
especially the bonding between composite and ceramic
and an analysis and rationalization of the results.
ACKNOWLEDGEMENTS
The authors would like to thank and acknowledge the
funding received from the Australian Dental Research
Foundation for their financial assistance given towards
the laboratory cost, the time and tireless energies of Ken
Tyler (University of Sydney) and Georges Sara (Stone
Glass Industries) for his help in fabricating the unique
inlay bridges required.
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Address for correspondence:
Dr Mark C Thompson
Faculty of Dentistry
The University of Sydney
Sydney NSW 2006
Email: mthompson@pacific.net.au
ª 2012 Australian Dental Association