The all-ceramic, inlay supported fixed partial denture. Part 2. Fixed partial denture design: a finite element analysis


Dr Mark C Thompson
Faculty of Dentistry
The University of Sydney
Sydney NSW 2006


The clinical use of all-ceramic crowns and fixed partial dentures has seen widespread adoption over the past few years due to their increasing durability and longevity. However, the application of inlays as an abutment design has not been as readily embraced because of their relatively high failure rates. With the use of an idealized inlay preparation design and prosthesis form which better distributes the tensile stresses, it is possible to utilize the inlay as support for an all-ceramic fixed partial denture. Utilizing a three-dimensional finite element analysis, a direct comparison of the inlay supported all-ceramic bridge against the traditional full crown supported all-ceramic bridge is made. The results demonstrate that peak stresses in the inlay bridge are around 20% higher than in the full crown supported bridge with von Mises peaking at about 730 MPa when subjected to theoretical average maximum bite force in the molar region of 700 N, which is similar to the ultimate tensile strengths of current zirconia based ceramics.

Abbreviations and acronyms:

computer-aided design/computer-aided milled


computer tomography


degrees of freedom


finite element analysis


fixed partial denture


Houndsfield units


lithium disilicate glass ceramic


periodontal ligament




yttria stabilized tetragonal zirconia


The idealized tooth preparation design for a ceramic inlay was described in Part 1 based on a comprehensive literature review.1 In this paper we review the literature regarding the all-ceramic fixed partial denture (FPD) design and analyse the inlay supported FPD via a finite element analysis (FEA) comparison vis-à-vis the conventional and proven full crown supported, all-ceramic FPD.

FPDs must conform to anatomical and physiological constraints which necessitates restrictions upon their dimensions and geometry. Proper designs for the pontic-abutment connector must relate favourably to the Euler–Bernoulli beam theory (or the simple law of beams) which simply stated implies that the deflection of a beam increases as the cube of its length, is inversely proportional to its width and is inversely proportional to the cube of its height.2 Ceramics used in place of cast metals as a substructural element in FPDs has seen the simple transference of conventional porcelain-fused-to-metal (PFM) designs to all-ceramic structures. Core ceramics and the newer monolithic ceramic systems possess significantly higher elastic moduli and strengths than ever before. Nevertheless, contrasts in elasticity and geometry issues largely determine stress distribution and hence modes of failure.3

The capacity of ceramics to bear loads is determined by their limited strength, low fracture toughness and associated high Young’s Modulus.4,5 This low fracture toughness is further impacted upon by the process of time-dependent, subcritical crack growth so markedly displayed by ceramic systems.6 These multitudes of microscopic surface flaws formed during processing lower the practical strength of ceramics by two to three orders of magnitude from the theoretical maximum of the perfect specimen and cause large variations in strength and a time dependency due to variations in the distribution and depth of the initial cracks and their time-dependent propagation to critical failure. This slow crack growth is the result of a chemical interaction between the ceramic or brittle solid and its environment, usually water which causes the hydration and hence weakening of the ceramics metal-oxygen-metal bond,7 and together with cyclic fatigue is an important factor in the failure of ceramics intraorally.8

Ceramic strength displays scatter due to the variability of flaw distribution which means that small samples are generally stronger than larger samples because of the reduced probability of finding large cracks in the material with less surface area. They are also generally stronger in bending than in tension due to less of the surface area of the sample being exposed to high uniform stress.7

Weibull9,10 developed the following formula to handle the statistics of fracture strength in recognition of the huge variability in strength and reliability of brittle materials


where P is the failure probability, σ1 the samples inert strength, σ0 the samples Weibull scaling stress and m the Weibull modulus which is a measure of the variability in strength of the material, corresponding to the shape of the distribution curve.5 A larger m relates to a more homogenous material and thus higher survival rates. Engineering ceramics typically have a Weibull variability of between 10 and 15 compared to 5 for window glass, hence demonstrating the greater reliability of these ceramics over domestic glass.7 Thus, complex materials such as ceramics require the use of FEA because of the difficulty in determining long-term survivability through bench-top testing alone.

Stresses, strains and shearing within the tooth/restoration complex is the result of numerous factors including the abutment preparation geometry (reviewed above), the morphology and geometric outline of the restorative/prosthodontic material,11 whether the material is homogenous or multi-layered,12 abutment conditions such as the material of the abutment (particularly its Young’s modulus) and whether the abutments are fixed or allowed to rotate under load. Clinically, the fracture resistance of ceramic FPDs is largely related to the size, shape and location of the connectors and stresses applied to the pontic span. The non-uniform, highly complex shape of dental prostheses and, in particular, the narrowing of the minor connector between abutment and pontic results in the concentration of stress. These stresses resolve themselves into compressive forces (at the occlusal surface) and tensile forces (at the gingival aspect) due to the relatively small radius of curvature at the embrasure.13

Literature review of the relationship between the gingival embrasure of an FPD and its long-term survival and methods used for testing

An extensive body of evidence has demonstrated or discussed that broadening the curvature of the gingival connector of FPDs results in better distribution of stresses.12–27 However, the relationship between radius of curvature and fracture resistance has not been examined in sufficient detail.13 The use of traditional load-to-failure bench-top testing is unable to recreate the failure mechanisms seen clinically,28 hence the use of FEA is gaining popularity because of its ability to accurately assess the complex biomechanical behaviour of irregular prosthetic structures and heterogeneous materials in a non-destructive, repeatable manner.

Kelly et al.18 conducted a fractographic analysis of 29 all-ceramic, crown supported FPDs which had failed; 20 after bench-top testing and 9 failing clinically. The coincidence of bench-top and clinical observations for the structural behaviour of these prostheses was examined. Detailed investigations into the in vitro prostheses proved that the failures had originated from the connector sites and in 70% of cases initiation was from the core-veneer interface. Most significantly, the cracks developed at the apex of the gingival embrasure and extended to the contact site. All 9 clinically failed FPDs shared this same mode of failure with the cracks originating from the gingival embrasure and extending to the site of loading.

A two-dimensional FEA conducted on the laboratory samples was consistent with the fractographic analysis, with peak tensile stresses localized around the connector area but only if there was a significant difference in the Young’s modulus between the core and veneering ceramics and, most interestingly, if a small amount of abutment rotation was allowed. The rigid fixation of the abutments displayed an FEA solution markedly different than one where the abutment was allowed to rotate or simply move due to the presence of the periodontal membrane, with the former FEA displaying a result akin to a classic four-point bend test, and the latter closely mimicking the fractographic analysis. Weibull calculations of fracture probability in both the FEA and fractographic analysis were very closely matched.

The major limitations of the above study were that the abutments were cast in Ni-Cr-Be, a material significantly stiffer than dentine and the absence of a simulated periodontal ligament, both of which led to the sample fracturing in a rather different manner than a dynamic biological structure. The elastic modulus of the supporting structure is very important in the distribution of stresses29–32 as demonstrated by the fracture resistance of all-ceramic crowns fixed on Co-Cr-Mo (E = 180–240 GPa) models being significantly higher than those cemented to natural teeth (E = 15–20 GPa for dentine and 50–85 GPA for enamel) – 1838 N vs. 888 N, respectively. Additionally, the connector dimensions and geometry were not carefully controlled. Nevertheless, the in vitro samples were able to closely mimic the results of the failed clinical models.

Fisher et al.14 utilized FEA to predict if computational methods could help FPD design improve in reliability. Their all-ceramic, crown supported FPD designs utilized three models, each differing in connector dimensions and connector length, roughly corresponding to embrasure radius, connector height and bucco-lingual depth as follows:


The results indicated that the Weibull curve moved significantly towards lower failure probability when the connector length increased as in the third design and cross-section increased from 8.75 mm2 to 14 mm2. Additionally, the FEA demonstrated that maximum stress decreased in the third group from 106 to 85 MPa. The authors reported that the effects of the periodontal ligament and cement film are negligible on the mechanics of fracture. However, other studies have been careful to account for these factors.

Kou et al.12 utilized a newly developed FEA model code to assess the fracture mechanism of yttria stabilized tetragonal zirconia (Y-TZP), crown supported FPDs. The study examined heterogenous Y-TZP under static loading and on stainless steel abutments, comparing it to a previous study by Sundh et al.27 which involved bench-top testing. Fracture patterns were very similar in both cases with all fractures initiating at the gingival embrasure, propagating diagonally under one single loading level (a property of brittle materials) and extending to the loading point. Additional analysis involving photo-elastic fringe patterns and acoustic emissions agreed with the results of the study by Sundh et al.27 and the current FEA.

The authors stated that a major limitation of their study was the simplified two-dimensional testing and like so many other studies, failed to account for the effects of a natural tooth’s modulus of elasticity, the influence of the cement film and the visco-elastic nature of the periodontal ligament. Ignoring the modulus of elasticity of the abutment and the luting cement has been shown to influence the fracture resistance of crowns33 and hence FPDs. The poor marginal fit of ceramic inlays and their lack of primary mechanical retention places greater importance on the role of the luting cement as compared to cast-metal inlays.34

Hojjatie et al.16 and Fischer et al.14 believed that the effects of the cement film and periodontal ligament could be ignored. However, Kelly et al.18 showed that ignoring the mobility provided for by the periodontal ligament restricts the rotational movement of the abutments, thus requiring greater fracture forces because the FPD functions more as a static beam rather than a dynamic anatomical structure. Likewise, Kappert et al.35 reported a mean fracture strength of 703 N for all-ceramic, three-unit FPDs cemented on abutments with simulated periodontal mobility compared to mean fracture strengths of 2225 N when fixed to ridged abutments.

Wolfart et al.36 bench-top tested a total of 64 all-ceramic, inlay retained FPDs made from heat pressed lithium disilicate glass ceramic (LDGC) (n = 32) and Y-TZP (n = 32) with half from each group made with different connector dimensions of 3 × 3 mm or 4 × 4 mm. A comparison of their quasi-static fracture resistance and fatigue strengths when changes were made to their connector dimensions in the two material groups was conducted. The abutments were Co-Cr-Mo, supported in an alloy base with simulated periodontal ligament made from silicone material and the FPDs were cemented with a composite luting agent.

The results for the median fracture strengths were as follows:


Statistically, the quasi-static loading and cyclic loading showed significant differences between both groups of LDGC (p ≤ 0.03) but not between the two groups of Y-TZP. The difference between LDGC and Y-TZP was significant for both connector sizes (p ≤ 0.001).

However, by using Co-Cr-Mo as the abutment material, this study failed to take into account the ability of dentine to redistribute stresses due to its significantly lower modulus of elasticity and, as demonstrated by Kappert et al.,35 can lead to fracture forces markedly different to that displayed by natural teeth. Moreover, the use of a Co-Cr-Mo may also not adequately bond to the composite resin luting cement, thus impeding the adhesion of the frameworks.

Oh and Anusavice13 tested the effect of various all-ceramic FPD connector designs based on the hypothesis that increasing the radius of curvature of the gingival embrasure leads to a decrease in fracture probability. The results concluded that variations in the occlusal embrasure radii was of no significance, with variations in gingival embrasure radii of between 0.25 mm and 0.90 mm increasing the mean fracture strength by more than 140% (p ≤ 0.0001).

Oh, Götzen and Anusavice22 further tested the hypothesis that increasing gingival radii curvature would lead to increasing fracture resistance of FPDs via 2 three-dimensional FEA models based on the original study by Oh and Anusavice.13 A change in the gingival radius from 0.9 to 0.45 mm was carried out in order to maintain a constant connector height of 4 mm. Fractographic analyses of the 40 failed FPDs from the Oh and Anusavice13 study were assessed and compared to the FEA, revealing that the failure origin was at the gingival embrasure in all 40 specimens. The results of the FEA correlated well with the fractographic findings, with peak compressive stresses occurring at the occlusal embrasure and peak tensile stresses at the gingival embrasure. Weibull moduli were 6.3 for the group with the narrow gingival embrasure and 8.6 for the group with the lager embrasure.

Plengsombut et al.37 studied the effects of a sharp (0.06 mm radius) vs. a round (0.6 mm radius) gingival embrasure shape on the fracture resistance of all-ceramic core materials, namely a pressed and milled LDGC and Y-TZP ceramic blocks. The results showed a significant difference in the fracture strength of the material due to, not only the gingival geometry, but also the fabrication technique. Specifically, if the material was machined, i.e. computer-aided design/computer-aided milled (CAD/CAM), then the connector design affected the fracture resistance. If on the other hand the material was pressed, then the connector shape had no significance on the material’s strength.

Lithium disilicate based materials possess larger flaws, a coarser surface finish compared to Y-TZP materials and hence were structurally weaker regardless of fabrication technique or connector shape. Y-TZP is inherently tougher and possesses a smoother surface finish; therefore is structurally strong and less influenced by connector shape, but still benefits from the effects of stress distribution afforded by the rounded gingival embrasure.

Magne et al.20 performed a two-dimensional FEA (the authors accepted that a three-dimensional FEA would more realistically model stresses and strains but at the cost of greater processing requirements) investigating: (1) the stresses at the surface and interface of 3-unit posterior adhesive FPDs made with composite resin, fibre-reinforced CR, gold alloy, LDGC, high alumina glass ceramic or Y-TZP; and (2) the influence of slot vs. 2-surface vs. 3-surface abutment preparations. The different materials and abutment preparations were constructed upon a digitized cross section of a 3-unit FPD. Included in the numerical analysis was the periodontal ligament (PDL) and supporting bone. A 50 N simulated vertical load was applied to the pontic and the stresses within the restorative materials. Tooth/restorative junction and surface stresses were calculated and compared.

The results concluded that for all materials, the stress patterns were remarkably similar to that displayed in a typical three-point bending test, with highest tensile stresses at the gingival surface (peaking at the gingival embrasure), compressive at the occlusal, and the abutments were subject to mainly compressive forces. Overall, the distribution of stresses was most favourable for the composite material due to its lower modulus of elasticity (12.3 GPa). However, this material was deemed to be clinically unsuccessful for FPD use because of its relatively low fracture toughness. With improving material toughness comes increasingly high stiffness and brittleness, culminating in the alumina ceramics and Y-TPZ with moduli of elasticity as high as 402 and 205 GPa, resulting in the placement of proportionately higher stresses on the interfacial surfaces and cement layers.

Kiliçarslan et al.38 tested the fracture resistance of 32 posterior, metal-ceramic crown supported (n = 8) and inlay-retained (n = 8) FPDs, and all-ceramic inlay retained FPDs (LDGC-n = 8, Y-TZP-n = 8) utilizing load-to-failure bench-top testing.

The results were that all metal-ceramic FPDs failed via adhesive failure of the ceramic-to-metal with 7 of 8 in the retainer area and 1 in the pontic. All-ceramic inlay supported FPDs failed via cohesive failure of the ceramic structure with 4 fracturing in the connector and 4 in the inlay retainer. The metal ceramic FPDs had the highest mean failure load at 1318 N for the crown supported group and 858 N for the inlay supported group. LDGC showed the lowest failure load at 303 N whilst the Y-TZP displayed static fracture strength of 1247 N, very close to that of the metal ceramic, crown supported group.

The use of metal models and the absence of a simulated PDL were accepted by the authors as a limitation. Further commentary must be made regarding the use of Sinogol (a provisional cement), thus lacking the proper adhesive qualities of a composite resin so needed in the bonding of the inlays, for the adhesive is paramount in the overall success of a bonded restoration.39 Brittle materials, such as ceramics, are weakest when exposed to tensile stresses.40 Hence, subjecting a material or design to such tensile stresses is considered to be an excellent test of its properties. Proper designs for the critical connector and pontic must relate favourably to elementary beam theory. Unlike beams, however, the dimensions and shape of FPDs is never uniform but heavily dependent upon the tooth preparation, which in turn is heavily influenced by the morphology of the remaining sound tooth structure after caries removal, the individual anatomical and geometric alignments of the abutment teeth, the length of the edentulous span and the luting agent used.

Summary of the literature conclusions

Based on our analysis of the literature, the following were concluded: (1) tensile stresses are concentrated at the gingival aspect of the connector and the vast majority of ceramic failures are initiated at this site; (2) increasing the gingival embrasure radii results in better distribution of stresses and higher fracture loads; (3) increasing the dimensions of the connector also results in higher fracture loads; (4) rigid abutment fixation results in fracture strengths and stresses different to abutments that are allowed to rotate via a PDL; (5) the elastic modulus of the abutment material is significant in accurately reproducing clinical performance; (6) the effects of a luting agent is important; and (7) three-dimensional FEAs are more accurate than two-dimensional.

Traditional load-to-failure testing methods have proved irrelevant in predicting the clinical performance of ceramics, largely because they cannot recreate the failure mechanisms seen in clinical specimens. Numerical analysis proves to be a more accurate predictor.41 With reference to the conclusions of the literature review above, it should be possible to construct an optimized all-ceramic, inlay supported FPD with stresses reduced to a minimum and test the design via an accurate numerical simulation viz. FEA.

This paper aims to test the design hypothesis that the use of an idealized inlay preparation geometry on the abutments and increasing the gingival embrasure radii and broadening the connector of the bridge, will minimize stresses within the all-ceramic inlay supported FPD to the degree where it will be clinically acceptable compared to the all-ceramic full crown supported FPD. Two three-dimensional finite element models were constructed for the biomechanical analysis and comparison. The connector designs were as follows: (1) inlay supported FPD – mesial connector height 3.5 mm, width 4.8 mm and gingival embrasure radius of 0.9 to 1.0 mm, distal connector height 3.5 mm, width 5.4 mm and gingival embrasure radius 1.4 mm average. These dimensions reflect the results of the above studies in minimizing the effects of stress and strain; and (2) full crown supported FPD – mesial connector height 4.0 mm, width 5.8 mm and gingival embrasure radius 0.8 to 0.9 mm, distal connector height 4.0 mm, width 6.0 mm and gingival embrasure radius 0.8 to 0.9 mm.

All material properties were kept standardized. However, in the interest of replicating clinical norms, the connector area of the crown supported prosthesis was deliberately enlarged and the embrasure radii reduced, thus taking into account clinical and laboratory norms in making use of as much connector height and width as possible. This bolsters the connector strength but reduces the gingival embrasure radii in order to improve interproximal aesthetics. Likewise the connector area on the inlay supported prosthesis was enlarged but due to the constraints of the inlay widths, the contact was necessarily smaller both in height and width.

Materials and Methods

Geometry acquisition

The FEA geometry was based on micro computer tomography (CT) imaging of a natural (cadaver) mandible section from the lower right posterior segment where the first molar was missing. The CT captured anatomical edges and these were digitized in Amira v 4.1.1 (Visage Imaging GmbH, Germany) to capture the geometry of the cortical bone, cancellous bone and dentine.

Geometry (bone and dentine) were then refined in Studio v9 (Geomagic Inc., USA). Geometries were imported into Rhinoceros 3D v 4.0 (McNeel Inc., USA), and then the PDL (0.3 mm) generated in the tooth sockets. The FPD was created utilizing the idealized inlay design described in Part 1 and full crown abutment utilizing 1 mm buccal and lingual, and slightly narrower interproximal marginal widths.


Three-dimensional mesh generation was developed with MSC Patran (Santa Ana, USA). Elastic moduli and Poisson’s ratio was entered for the ceramic, PDL, dentine, cortical and cancellous bone. Houndsfeld units (HU) from the CT data were used for the cortical and cancellous bone values and densities were created. These densities were in the range of 700 HU for cancellous and 1700 HU for cortical bone.42,43

A static normal load of 200 N was applied to the centre of the first molar abutment, corresponding to a steel ball of approximately 5 mm2; 200 N was chosen because it corresponds closely to the maximum intercuspation force on the mandibular molars,44 whilst allowing us to easily scale the force due to the linear-elastic or Hookean behaviour of ceramics.

The FPD-abutment interfaces were tied to simulate full cement bonding. Edges of the mandibular bone were fixed. Linear elastic, homogenous material properties were assumed for the PDL, abutments and FPD45,46 as summarized in Table 1. Heterogeneous material properties were applied to the cortical and cancellous bone structures. It was assumed during normal mastication function the PDL was loaded in its elastic range and its behaviour could be defined approximately linearly.

Table 1.   Material properties required within the FEA models47
MaterialYoung’s Modulus (MPa) Poisson’s Ratio
Enamel84 1000.20
Dentine18 6000.31
FPD ceramic200 0000.28

The resulting 2 three-dimensional FEMs (inlay and full crown abutments) were the basis for the biomechanical analysis. These models were meshed using 10-node quadratic tetrahedral elements with a global element size of 1 mm in MSC/PATRAN and consisted of 117 220 elements and degrees of freedom (DOF) 724 461 and 113 552 elements and DOF 697 713, respectively.


Resultant geometry was brought into the FEA programme ABAQUS 6.6.1 (Dassault Systèmes, France) for post-processing, with the analysis displaying principal and von Mises stresses.


Figures 1 and 2 display the resultant von Mises (also known as the Distortion Energy Theory) stress contours. These clearly show that the highest stress concentrations exist in the vicinity of the embrasure areas between the inlay and pontic and at the loading contact site, peaking at 209 MPa in the inlay model and 174 MPa in the crown model. Von Mises stress output does not define these stresses as tensile (negative) or compressive (positive) but only describes their magnitude. Determining whether the stresses are tension or compression require an examination of the principle stress values of the object. Also notable is the increased degree of stress throughout the body of the inlay (as noted by the hot spots), especially at its axio-pulpal line angle as opposed to the body of the crown.

Figure 1.

 von Mises stress contours for the inlay supported FPD.

Figure 2.

 von Mises stress contours for the full crown supported FPD.

Figures 3 and 4 reveal the resultant FEA maximum principle stress contours. High tensile stresses are evident at the gingival aspect of the connectors peaking at 198 MPa in the inlay bridge and 177 MPa in the full crown bridge.

Figure 3.

 Maximum principle stress contours for the inlay supported FPD.

Figure 4.

 Maximum principle stresses for the inlay supported FPD.

Figures 5 and 6 reveal the resultant FEA minimum principle stress contours. High compressive stresses are evident at the gingival aspect of the connectors peaking at 177 MPa in the inlay bridge and 128 MPa in the full crown bridge.

Figure 5.

 Minimum principle stress contours for the inlay supported FPD.

Figure 6.

 Minimum principle stress contours for the full crown supported FPD.


In order to understand the consequences of the results of the stresses and strains within the ceramic structures, it is necessary to compare and analyse the data in terms of theories of failure which best describe events. Comparisons were initially made in terms of von Mises criteria and principal stresses. Firstly, von Mises is a formula for combining the three principal uniaxial stresses (x, y and z planes) into an equivalent ‘scalar’ applied stress. It is one of several failure theories used to determine the applied uniaxial stress in a structure and provides a measure of the degree of overall stress. Most importantly, it takes into consideration deviatoric energy or shear energy as a result of an applied load. It is most relevant for yield or plastic flow induced failure as with metallic materials. Our understanding of the physical world is made via direct comparisons from simple uniaxial bench-top stress and strain testing. However, this correlates poorly with actual failure in a triaxial state. If an elastic body is subject to a system of three-dimensional loads, what we typically find is that even though none of the uni-axial principal stresses exceed the yield stress of the material, it is possible for yielding to occur from the resultant combination or the three stresses. The equivalent stress, called the ‘von Mises stress’, if exceeding the yield stress of the material results in ductile failure.

Secondly, the maximum and minimum principal stresses were evaluated to accurately disclose areas of tension and compression. Principle stress is considered most useful for brittle materials and states that failure will occur when maximum principal stress developed in a body exceeds the uniaxial ultimate tensile/compressive strength of the material. It is generally agreed that compressive hydrostatic (volumetric) pressure does not ultimately lead to material failure, shearing and energy theories are relevant to ductile yield failure while principal tensile stresses are most critical for brittle materials failure.48

Both failure theories show that the overall distribution and peak values for stress favours the full crown supported bridge with the contours of stress clearly being more gradual and the peak stresses being not only lower but more confined than seen in the inlay supported bridge. The occlusal surfaces displayed a general compressive stress pattern whilst the gingival surfaces of the pontics displayed tensile stresses (Figs 3 and 4), peaking at the gingival aspect of the connector area for both designs. Additional tensile peaks are observed at the axio-pulpal line angle of the inlays (Figs 1c and 1d). Both designs displayed the classic stress pattern distribution of a three-point bending test as indicated by Magne,20 with the lower two points being located in the vicinity of the inter-proximal margins of the abutments.

The full crown supported bridge displays a more favourable stress distribution pattern largely due to the added bulk material available being better able to absorb and distribute stress, the Vierendeel truss or open torsion box-like nature of the full crowns and the added support gained from the margins of the crowns. This adds considerably to the rigidity of the structure, leading to less deflection and thus more even stress distribution.

von Mises stresses were 20% higher for the inlay bridge than the full crown bridge whereas the maximum and minimum principle stresses predict that the tensile stresses in inlay connectors were 12% and 38% greater, respectively compared to the crown connectors. The actual values in the connector areas would depend directly on the maximum load applied. In the present simulation a load of only 200 N was chosen. If higher loads were experienced, the maximum value would be expected to scale directly with the increasing load as the modelling conducted here was an elastic analysis. Hence, with maximum bite force in the molar region usually reported between 600 to 750 N,49 we merely need scale the stresses in the above FEM by a factor of three and a half to find the equivalent von Mises (730 MPa for the inlay bridge vs. 610 MPa for the full crown bridge) and the maximum tensile stresses (693 MPa for the inlay bridge vs. 620 MPa for the full crown bridge) under theoretical maximum oral loading conditions. Both are well within the reported ultimate fracture strength of Y-TZP of between 900 and 1200 MPa.50 On the basis of the current FEA, it would be expected that a Y-TZP inlay supported FPD would have an ultimate load of up to 1150 N according to the von Mises theory and 1240 N according to the Maximum Principle Stress Theory. This compares to 1380 and 1360, respectively for the full crown supported FPD; both reported nearly double maximum bite force.

The clinical use of ceramics and especially Y-TZP for the fabrication of full crown supported FPDs has become common place in clinical dentistry due to their natural aesthetics and exceptional biocompatibility. However, failure rates have proven to be greater than traditional metal-ceramic structures but low enough to be acceptable based on current popularity of all-ceramic systems. The inlay supported all-ceramic FPD is not considered routine or common place due to the perception that the incidence of failure relating to its use is insurmountable. However, the results of the FEMs detailed above indicate stress increases of only 20% greater for the inlay model over the full crown model and being able to withstand a theoretical ultimate yield load of nearly 1200 N, but a realistic load closer to 700 N, Y-TZP should be considered capable of clinical use as a material in the fabrication of inlay supported FPDs. Other ceramics such as the glass-infiltrated ceramics and densely sintered alumina, with fracture strengths of between 600 and 700 MPa may be considered for use, but with lesser safety margins.


The literature review presented earlier has enabled us to develop a more optimized design for the fabrication of an all-ceramic inlay supported bridge and numerically test the design against the accepted clinical norm of the full crown supported FPD. This testing utilized the current state-of-the-art three-dimensional FEA method, incorporating the conclusions and findings from previous studies, the use of which entails a non-destructive and repeatable way to test the performance of materials and designs.

The FEM presented above demonstrates that with the use of an idealized inlay preparation form (as discussed in Part 1)1 and an optimized bridge design emphasizing a broadening of the gingival embrasure, the forces derived from mastication can be adequately distributed to levels which are within the fracture strength of current ceramics.