Mehrali Et Al. [2013] Dental Implants From Functionally Graded Materials

Review Article

Dental implants from functionally graded materials

Mehdi Mehrali,1 Farid Seyed Shirazi,1 Mohammad Mehrali,2 Hendrik Simon Cornelis Metselaar,2

Nahrizul Adib Bin Kadri,1 Noor Azuan Abu Osman1

1Department of Biomedical Engineering, Faculty of Engineering, University of Malaya, Kuala Lumpur 50603, Malaysia2Department of Mechanical Engineering and Center of advanced Material, University of Malaya, Kuala Lumpur 50603,

Malaysia

Received 29 August 2012; accepted 4 January 2013

Published online 11 June 2013 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.34588

Abstract: Functionally graded material (FGM) is a heteroge-

neous composite material including a number of constitu-

ents that exhibit a compositional gradient from one surface

of the material to the other subsequently, resulting in a ma-

terial with continuously varying properties in the thickness

direction. FGMs are gaining attention for biomedical applica-

tions, especially for implants, owing to their reported supe-

rior composition. Dental implants can be functionally graded

to create an optimized mechanical behavior and achieve the

intended biocompatibility and osseointegration improve-

ment. This review presents a comprehensive summary of

biomaterials and manufacturing techniques researchers

employ throughout the world. Generally, FGM and FGM po-

rous biomaterials are more difficult to fabricate than uni-

form or homogenous biomaterials. Therefore, our

discussion is intended to give the readers about successful

and obstacles fabrication of FGM and porous FGM in dental

implants that will bring state-of-the-art technology to the

bedside and develop quality of life and present standards of

care. VC 2013 Wiley Periodicals, Inc. J Biomed Mater Res Part A:

101A: 3046–3057, 2013.

Key Words: functionally graded material (FGM), biomaterials,

dental implants, mechanical properties

How to cite this article: Mehrali M, Shirazi FS, Mehrali M, Metselaar HSC, Kadri NAB, Osman NAA. 2013. Dental implants fromfunctionally graded materials. J Biomed Mater Res Part A 2013:101A:3046–3057.

INTRODUCTION

The aim of modern dentistry is to restore the patient tonormal function, health, speech, and aesthetics. The dentalimplants could be an only restoration option for people ingood general oral health who have lost a tooth owing toperiodontal disease, an injury, or some other reasons. Themost common cause for failure in dental implant is inad-equate bone formation around the biomaterial immediatelyafter implantation.1 Therefore, the development of new bio-materials for dental implants is one of the challenging tasksfor materials science today. A single composition with a uni-form structure cannot satisfy the requirements for somebiomedical applications as dental biomaterials must meetseveral important criteria and have properties such as bio-compatibility with a known degradation rate, osteoconduc-tivity, strength, corrosion resistance, elastic modulus, fatiguedurability, and close chemical similarity to biological apatitepresent in human hard tissues.2–5 Hence, in the field of bio-medical implants, we often observe a number of designs

produced both from conventional engineering and tissue en-gineering. Nowadays, bone implant research is mainlyfocused on four areas: (1) composites, (2) polymeric coat-ings on metallic implants, (3) tissue engineering, andrecently (4), functionally graded material (FGM).6

FGMs, first proposed in 1986 in Japan, represent a novelidea for the realization of innovative properties and/or func-tions that conventional homogeneous materials cannot ac-complish.7,8 One extraordinary feature of biomaterials is theformation of gradable structures. Therefore, FGMs are usefulas the composition of tissue shows a continuous changefrom one composition to another.9 For example, the suitabledesign of porous bone with a porosity gradient from adense, stiff external structure (the cortical bone) to a porousinternal one (the cancellous bone), and with an adequatedegree of interconnectivity exhibits that functional gradationis applied by biological adaptation.10–12

In the case of dental implants, the components are usu-ally much smaller and used to reconstruct the masticatory

Correspondence to: F. S. Shirazi; e-mail: [email protected]

Contract grant sponsor: Ministry of Higher Education (MOHE) of Malaysia; contract grant number: UM.C/HIR/MOHE/ENG/10 D000010-16001

Contract grant sponsor: Institute of Research Management and Consultancy (IPPP), University of Malaya; contract grant number: PV008/2012A

3046 VC 2013 WILEY PERIODICALS, INC.

function when a tooth root is wholly extracted or lost. Theimplant is located in the jaw bone in a manner to penetratefrom the inside to the outside of the bone. The requiredfunction of a dental implant varies at the outside of the jaw-bone, inside it, and at jawbone boundary.13 On the outsideof the bone, the implant material needs to have sufficientmechanical strength to bear the occlusal force, whereas thepart inside the jawbone must have stress relaxation, osteo-conductivity, and adequate bone-implant contact so that thenew bone is created speedily and attaches directly to it.Nowadays, dental implants use single materials sometimeswith a bioactive coating layer. The problems associated withthe current, coated metallic implants consist of stressshielding of the surrounding bone and poor survival of coat-ings over time, resulting in severe biocompatibilityissues.6,14 The theory of stress shielding hypothesizes thatbone loss around orthopedic and dental implants is owingto the removal of normal stress from the bone by animplant.15 To improve the dental implant’s acceptanceinside the living bone, the concept of FGM should be afavorable approach.16,17 The primary advantages of usingFGM dental implants include: (1) improvement of biocom-patibility,18 (2) diminution of the stress shielding effect onthe surrounding bones that regularly occurs in the presenceof fully metallic implants,19–21; (3) precluding the thermalmechanical failure at the interface of hydroxyapatite (HA)-coated metallic implants,22 and (4) improving the biome-chanical requirement and controllability for graded bioreac-tion at each region of the bone.23 All these aspects of usingFGM implants are discussed in this review. So far, function-ally graded dental implants use a variety of materials suchas titanium (Ti), HA, titanium nitride (TiN), polymer, zirco-nia (zirconium dioxide, ZrO2), and so on. In this review arti-cle, we consider materials and fabrication techniques usedin this application field in recent years. Furthermore, the

design of functionally graded dental implants and optimiza-tion is also a subject of this review.

THE STRUCTURE OF FUNCTIONALLY GRADED DENTAL

IMPLANTS

Many natural tissues and organs are not homogeneousmaterials, and observations show their structures to befunctional gradients, for example mollusk shells, bamboo,bone, and skin.24–26 Tissues or organs are described asfunctionally graded if each layer of the tissue or organ hasone or more particular functions to achieve the local func-tional requirements. Therefore, to regenerate the naturalfunctionality, a successful conception of dental implantshould also include FGM.5,17,27–30 Figure 1 shows a sche-matic view of an FGM dental implant with graded materialcomposition used in dentistry.

For FGM dental implant purposes, a cylindrical shapewas designed with composition varying in the axial direc-tion. The upper part has more strength and mechanicalproperties necessary as the occlusal force is applied directlyon the top and then transmitted down to lower partsimplanted inside the trabecular bone where more biocom-patible materials are desirable.6,31,32 Mostly, implant failureis owing to the lack of biomechanical bonding between theimplant and the surrounding jawbone; therefore, the lowerparts should have a high osteoconductivity and have a goodbone-to-implant contact for regenerating bone and rapidosseointegration.20,33–35 Also, the implants might fail owingto insufficient strength and mechanical properties or over-load. Consequently, a possible solution is to design adequatemechanical properties in the upper parts of FGM dentalimplants.34–37 As discussed in detail above, when consider-ing an implant design, the use of FGM is expected toimprove both the mechanical and the biological perform-ance of dental implants.

FIGURE 1. View of an FGM dental implant with graded material composition in the maxilla.13 [Color figure can be viewed in the online issue,

which is available at wileyonlinelibrary.com.]

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JOURNAL OF BIOMEDICAL MATERIALS RESEARCH A | OCT 2013 VOL 101A, ISSUE 10 3047

CHALLENGES FOR MATERIALS

In the last few years, a variety of biomaterials and biocer-amics were investigated for FGM dental implants such asTi/HA,16,17,38 titanium/cobalt (Ti/Co),39 Ti/ZrO2,

40 tita-nium/silica (Ti/SiO2),

41 and TiN/HA.42,43 Table I summa-rizes the Young’s modulus, Poisson’s ratio, mass density ofbone, and material densities for reference.19,31,44–53 Studieshave shown that Ti and its alloys are among the most suc-cessful metallic biomaterials for dental and orthopedicapplications because of their good mechanical properties(elastic modulus, toughness and fatigue, strength), excellentcorrosion resistance, and good biocompatibility.58–60 How-ever, Ti and Ti alloys are bioinert and cannot promote tissuebonding to the implants. Generally, a dental implant shouldhave sufficient mechanical strength to maintain integrity;for this reason, Ti and Ti alloys were used along the longitu-dinal direction from Ti, which is rich in the upper partswhere occlusal force is directly applied to ceramics or met-als rich in the other material that is implanted inside thejawbone.

Watari et al.39 investigated Ti/Co FGMs. They evaluatedthe mechanical properties and biocompatibility for implants.They reported that in the implantation analysis of Ti/Cointo the soft tissue, a thin fibrous connective tissue wasformed in the pure Ti region while the thickness of the fi-brous tissue layer increased with the Co concentration, lead-ing to inflammation problems at the Co-rich part. Therefore,the CO’s change of concentration seemed to affect the bio-compatibility of implant material. Other clinical studies alsoindicated that epithelial cells as well as fibroblasts have astronger negative response to a Co-chrome alloy than toTi.61 In this context, Co exhibited several advantages in me-chanical properties, wear resistance, and good corrosionresistance.

Takahashi40 studied the use of Ti/ZrO2 FGM for dentalimplants. Zirconia appears to be an appropriate dentalimplant material because of its tooth-like color, biocompati-bility, and good mechanical properties in strength and frac-ture toughness.62 The inflammatory reaction and boneresorption provoked by zirconia particles are less thanthose influenced by Ti particles; therefore, ZrO2 is suitable

for long-term use in vivo.63–65 Recently, Fujii et al.66,67

described the successful preparation of partially stabilizedzirconia (PSZ) and pure Ti FGM by the hot pressing of pow-der. Fujii et al. evaluated the mechanical properties such asbending strength, Young’s modulus, and Vickers hardness.One of the results was that the bending strength and Vick-ers hardness decreased with increasing Ti content in thePSZ–Ti because the strength and hardness of 100% PSZ aremuch higher than those of pure Ti. The Young’s modulusincreases from pure PSZ to 20% Ti content and thendeclined with increasing Ti. Also, they found that brittlefracture occurred after elastic deformation except for pureTi.66 Figure 2 shows the fracture surfaces in SEM micro-graphs. They reveal that transgranular fractures occurred inthe PSZ and Ti phases. It is worth noting the brittle fracturesurface is detected even in the Ti-rich region.

Takahashi et al.41 carried out the synthesis of Ti/SiO2

FGM for dental work in 1992. Some studies reported thatsilica is believed to play a critical role in bioactivity of bio-active materials for the bonding of bone and muscle and asa crosslinking agent in connective tissue.68–70 There arethree functions of silica. (1) Silica performs a specific meta-bolic function that is thought to partake in cellular develop-ment and gene expression.71–73 (2) There is a chemicalfunction in which the bonding to bone is established by theprecipitation of apatite surface layer that must be formedon the bioactive silica-based glasses surface when in contactwith in vivo applications or simulated body fluid.71,74,75 Thein vitro studies observed that composites containing bioac-tive glass–ceramic (BGC) nanoparticles with lower phospho-rous and higher silica content have better bioactivity thanthat of the BGC with higher phosphorous and lower silicacontent.76,77 (3) There is a mechanical function as silica par-ticles showed to improve the strength of a HA coating byparticle-mediated reinforcement, leading to crack deflectionor crack arrest. Amorphous SiO2 is a good candidate toenhance the mechanical properties of HA coatings.68

HA is one of the best among the bioactive materials as itcan bond to human bone because osteoblast cells penetrateinto HA. Therefore, HA is clinically applied to the teeth.78–80

However, the mechanical strength of HA is remarkably low

TABLE I. Materials and Bone Properties

MaterialsYoung’s

Modulus (GPa)Poisson’s

RatioDensity(kg/m3) Reference

Cortical bone 14 0.3 1700 Rho et al.44

Cancellous bone 3 0.3 270 Rho et al.44

Enamel 10.5 0.33 2906 Dowker et al.45 and Kitagawa et al.46

Dentin 19.7 0.33 1800 Kitagawa et al.47 and Angker et al.48

Ti 110 0.35 4500 Benzing et al.49 and Lin et al.50

HA 40 0.27 3219 Hedia and Mahmoud19

TiN 251 0.29 5300 Stone et al.51, Sherif El-Eskandarany et al.52,and Namazu et al.53

Co 210 0.31 8900 Marti54

ZrO2 200 0.31 6000 Gahlert et al.55 and Andreiotelli and Kohal56

SiO2 75 0.17 2200 Zhou et al.57

3048 MEHRALI ET AL. DENTAL IMPLANTS FROM FGM

and the application of HA to teeth was limited.81–83 Anotherstudy by Watari et al.16,17 used HA and Ti to prepare FGMdental implants. This team reported that a good combina-tion of HA and the mechanical properties of Ti is consideredas a promising approach to fabricate suitable FGM dentalimplants. During sintering of Ti/HA FGM, the HA is notdecomposed at 850�C.42 On the other hand, this tempera-ture is too low for sintering HA and mechanical propertiesdid not develop. Kondo et al.42 used TiN to improve the sin-tering of HA at higher temperature. They reported that byusing TiN instead of Ti, the decomposition of HA can besuppressed. However, the sintering of the TiN-rich regionwas still inadequate for the temperature up to 1200�C.

To avoid these problems, researchers have used differentfabrication techniques. In the following sections, there is areview of the processing methods to fabricate FGM dentalimplants. As mentioned earlier, PSZ has good biologicalcompatibility and a high mechanical strength; however, it isnot expected that PSZ embedded as an artificial bone com-bines with a bone because it has no osteoconductivity.When a material devoid of osteoconductivity is applied forbone implantation, there is a gap between the material andthe bone that grows with the passing of time and causesboth pain and abrasion of the material.

Many have carried out several investigations to solvethis problem by producing a composite of HA and PSZ.84–86

Matsuno et al.87 reported that a laminated HA/PSZ-sinteredcomposite with a gradient composition can be producedfrom HA and PSZ. This study said to expect wide receptionof HA/PSZ gradient composition as a biomaterial for hardtissue because it has osteoconductivity and high mechanicalstrength. Guo et al.88 have used yttria-stabilized tetragonalzirconia (Y-TZP) to develop functionally graded HA/ZrO2

composites. The authors reported that HA/Y-TZP function-ally graded composites showed considerable improvementin mechanical properties, whereas the HA phase in the com-posite layers was stable up to 1200�C and the Y-TZP secondphase remained the tetragonal zirconia (t-ZrO2) phase afterprocessing at the highest temperature of 1250�C. Figure 3 isa back-scattering electron (BSE) micrograph of a cross-sec-tion of the HA/Y-TZP FGM. The microhardness and theYoung’s modulus increased stepwise from the pure HA layer

to the HA þ 40 wt % Y-TZP layer across the HA/Y-TZP func-tionally graded composites as shown in Figure 4(a,b).

Besides the FGM dental implants discussed above, po-rous functionally graded biomaterial such as Ti and itsalloys also work. Nowadays, most oral implants are fabri-cated from Ti-6Al-4V (90% Ti, 6% aluminum, and 4% vana-dium) and pure Ti. Although Ti and its alloys have goodphysical, mechanical properties, corrosion resistance, andbiocompatibility,89,90 the stiffness of metallic dental implantsis not well-adapted to bone. This situation leads to stressshielding from the residual bone, which may result in detri-mental resorptive bone remodeling.91 The stiffness of ametal is determined by the Young’s modulus of the materialused as well as its area of moment inertia. In the case of Tidental implants, the Young’s modulus is far higher than thatof cortical bone as summarized in Table I.44 Therefore, it isproposed that the level of porosity is gradable, from a moreporous surface layer to a denser core, giving the potentialto have the same stiffness of the bone tissue at theimplant–bone interface.92,93

Hirschhorn94 in the 1970s used the concept of function-ally graded porous biomaterials to fabricate femoral stems.However, the study appeared abandoned owing to the

FIGURE 3. BSE image of the HA/Y-TZP functionally graded

composite.88

FIGURE 2. SEM micrographs of fracture surface of PSZ-Ti.66

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concerns over poor fatigue performance until a resurgenceof interest in 1995 by Becker et al.95 Becker investigatedthe mechanical properties of different biomedicallyapproved alloys such as 316L stainless-steel, Co-29Cr-6Moalloy, and Ti-6Al-4V alloy.95 Becker and Bolton’s study92

showed that it is possible to use porous FGM Ti alloys fororthodontic or maxillofacial implants.

It is possible to alter and optimize mechanical propertiesof porous materials in dental implants by controlling poros-ity, pore size, and pore distribution. Oh et al.96 investigatedthe effects of different pore sizes in vitro and in vivo andfound a suitable size range of 5–15 lm for fibroblastingrowth, 70–120 lm for chondrocyte ingrowth, and 100–400 lm for bone regeneration, depending on the porosity aswell as scaffold materials. Meanwhile, the surface microstruc-ture of dental implants is another crucial design feature.Some have reported that rough surfaces can promote betterand faster bone apposition as they are more osteoconductivethan smooth surfaces.97 Wong et al.98 and Carlsson et al.99

reported that a surface with Ra between 1 and 10 lm wasoptimal for promoting bone apposition and stimulating mes-enchymal cell differentiation into functional osteoblasts.

The following section outlines the various fabricationmethods of porous FGM for dental implants and providesan indication as to how far they can provide control overthese parameters. Different synthesis methods for FGM den-tal implants are focused in the next section.

FABRICATION TECHNIQUES AND MECHANICAL BEHAVIOR

FOR FUNCTIONALLY GRADED DENTAL IMPLANTS

In this section, we discuss the processing methods to fabri-cate FGM dental implants and their mechanical behavior.There are several methods for the fabrication of FGM andporous FGM dental implants as summarized in Table II.

Functionally graded dental implants containingTi and TiNTi/HA FGM can make a promising material for tissue im-plantation, orthopedic, and dental applications because ofits outstanding biocompatibility and bioactivity. As men-tioned above, different fabrication methods are used such ascold isostatic pressing (CIP), spark plasma sintering (SPS),hot pressing, and powder metallurgy (PM). Watari et al.17 in1997 showed the synthesis of HA/Ti as an FGM dental

FIGURE 4. (a) Microhardness and (b) Young’s modulus as a function of Y-TZP content for the HA/Y-TZP functionally graded composites by

SPS.88

TABLE II. Materials and Fabrication Techniques

Functionally GradedDental Implants Fabrication Methods References

Ti/HA CIPþEF, CIPþHF, SPS,PM, hot pressing

Watari et al.16, Fujii et al.66, Watari et al.100,Chu et al.101, and Chenglin et al.102

TiN/HA SPS Kondo et al.42

Ti/Co PM Watari et al.39

Ti/SiO2 CIPþHF Watari et al.16 and Takahashi et al.41

Ti/ZrO2 CIPþHF, hot pressing Takahashi40 and Fujii et al.66

HA/ZrO2 SPS, hot pressing Guo et al.88

Porous FGM PM, one-step microwave,PECS, LENSTM, and DLMS

Matsuno et al.87, Traini et al.93, Kutty and Bhaduri103,Suk et al.104, and Krishna et al.105


implant using CIP and sintered by high-frequency inductionheating to satisfy both mechanical and biocompatible prop-erties. Watari et al.16,38,100 tried CIP, sintered electric fur-nace (EF) heating and SPS methods for sintering. In the EFsintering, powders were packed into the thermo-contractivetube after heat treatment of a tube at 60�C, which was thencompressed by CIP at 800–1000 MPa, and the implants ofthe miniature cylindrical shape were densified by sinteringin a vacuum at 1300�C. In the high-frequency inductionheating, they packed powders into the thermo-contractivetube or silicone rubber impression mold with the shape of adental implant. After CIP, the packing was sintered at above1300�C in Ar gas atmosphere. For SPS process, the mixedpowders of Ti hydrate and HA were put into a graphitemold with the gradient composition ranging from 100% Tito pure HA in the height direction and sintered at 850�C ata pressure of 40 or 80 MPa. Watari et al. reported that Ti/100HA FGM decomposition in the sintering process is sig-nificant. Internal stress also arises from the difference ofthermal expansion coefficient and shrinkage at the interfacefrom one region to other. Moreover, they lowered the sinter-ing temperature by using SPS to avoid autodestruction.There was much improvement in sintering by SPS comparedwith the conventional CIP and furnace sintering methods.

Fracture of the FGM occurs near the weakest region orits neighbor. In the three-point flexural test of Ti/HA FGMprepared with SPS at a pressure of 40 MPa, fractureoccurred deviated from the center in the HA-rich side,which is weaker. In the FGM prepared with 80 MPa, fractureoccurred inside the single layer in the center. The flexuralstrength was increased to 36 MPa and the compressivestrength 88 MPa, respectively. The Brinell hardness testobserved for FGM sintered by SPS and EF in Figure 5 shows

the decreasing tendency of the hardness in the direction ofthe tooth root region in SPS contributes to the stress relaxa-tion, which relieves the jawbone from damage by impositionof high spike of impact stress near the implant.

It is important to control the temperature distribution ina work piece during and after sintering in SPS and the re-sidual stress distribution at the interface of each layer aftersintering to obtain the desired mechanical properties. Sasakiand Asaoka106 investigated and analyzed the effects of tem-perature distribution and residual stress at the interface ofeach layer in SPS sintering by finite element method (FEM)to develop an optimization approach for conducting opti-mum production conditions. FEM analysis can be used withSPS method to simulate the whole heating process in amanner comparable to the actual experimental experience,where a controllable heating rate, variable die size, andpressure are implemented.107,108 They reported that themeasured residual stress was higher than that predictedfrom FEM. This group also observed that hardnessincreased, whereas fracture toughness and bending strengthdecreased with rising HA content.

Chu et al.109 investigated HA/Ti biomaterial FGM byemploying a hot-pressing method without bending deforma-tion and microcracks parallel to the graded direction of theFGM on the surfaces. Chu et al.101 developed HA/Ti FGM byoptimizing the best combination of their biocompatibilityand mechanical properties. In their method, they first exam-ined the thermoelastic properties of uniform HA/Ti relatedto each graded layer of the FGM. In their results, the ther-mal expansion coefficients of the HA/Ti increased with therise in temperature or content of HA, and the residual ther-mal stress was confirmed by an X-ray and theoreticalmethod. They have also used powder metallurgy to produceHA/Ti FGM sintered at 1100�C.102 In this method, hardnessgrew with increasing volume fraction of HA as opposed toSPS sintering in which the hardness decreased. It is impor-tant to note that the Young’s modulus is higher than naturalbones in all regions of Ti/HA FGM that can cause severestress concentration, namely load shielding from a naturalbone, which may weaken the bone and deteriorate theimplant/bone interface. Chu et al. reported that the exis-tence of Ti can promote decomposition of HA at the sinter-ing temperature (1100�C). The decomposed phases are a-Ca3 (PO4)2 and Ca4P2O9. However, no new compounds formbetween HA and Ti.102

Kondo et al.42 studied TiN/HA functionally gradedimplants by SPS method. They reported that by using TiNinstead of Ti and sintering at 1100 and 1200�C, the decom-position of HA is suppressed. It was found that the mechani-cal properties TiN/HA are comparable to Ti/HA obtained bySPS and are sufficient for practical use. The Brinell hardnesswas around 60, whereas the hardness in the HA region didnot decrease. The flexural strength of TiN/HA sintered at1100 and 1200�C was 65.4 and 71.3 MPa, respectively, andthe compressive strength of TiN/HA FGM was more than100 MPa. However, the sintering of the TiN-rich region wasstill insufficient even at 1200�C. They observed little inflam-matory reaction at the TiN part in animal experiments.

FIGURE 5. Change of Brinell hardness with HA content in Ti/HA FGM

sintered by EF at 1300�C and SPS at 850�C.16

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Watari et al.39 produced Ti/Co FGM by powder metal-lurgy using both a wet technique to produce a gradient bydifferentiated sedimentation in a solvent liquid and a drymethod to pack mixed powders gradiently into a mold, fol-lowed by CIP, and sintering at 1300�C. The Ti/Co FGMspecimens were implanted in the subcutaneous tissue ofdorsal part of rat and after implantation were evaluatedfrom histological observation by optical microscopy. One ofthe evidences of infection is pus formation that is a resultof macrophages and neutrophils, which die after killing theforeign parasites. To avoid the spread of infection, fibroustissues usually form around the pus.110 The thickness andtexture of fibrous tissues created around the implant candepend on the type of implant material, shape and size ofthe implant, site of surgery in terms of functionality, andtype of tissue that needs to be healed.111 Watari et al.39

have found that the thickness of the fibrous connective layeris small in the pure Ti region. However, the thickness of thefibrous connective layer increased with Co concentrationalong the longitudinal direction of the FGM. It was alsoobserved that more inflammatory tissue and necrosisoccurred near some Co-rich parts.

Takahashi et al.41 manufactured Ti/SiO2 FGM dentalimplants by CIP method and then sintered in argon gas byhigh-frequency induction heating at a temperature of1300�C. This group reported that the fracture stresschanged from 2000 MPa in pure Ti to 100 MPa at the ce-ramic content. In their research, they also used Ti/ZrO2

FGM for dental implants40 by following the same method asmentioned above for Ti/SiO2. It was found that at the samevolumetric content, the Ti–zirconia composite showed thesame flexural strength as the Ti–apatite composite andhigher values of elastic modulus and strain. It was shownthat the sintering at higher ceramic contents could improveby adding 5% of Pd to the Ti.

Recently, Fujii et al.66 developed Ti and PSZ FGM usinghot pressing. The Young’s modulus and Vickers hardnesswere higher in the full range of Ti than predicted from therule of mixture, and the bending strength decreased withincreasing Ti content as the strength of pure PSZ is muchhigher than that of pure Ti. According to the X-ray diffrac-tion (XRD) analysis, the sintering process created reactionproducts such as Ti oxide, and these reactions influencedthe mechanical properties.66 It is often stated that there aredifferent stoichiometrics in Ti oxide which are TiO2, TiO,Ti2O3, Ti3O2, Ti3O5, TiO0.325, and TiO0.5, and it seems thatthe Young’s modulus depends on the stoichiometrics.112

Hence, various Young’s modulus of Ti oxide have beenreported.67,113 Hence, the mechanical properties of the Ti/PSZ FGM could not be explained by the rule of mixture.Teng et al.114 reported that there was no reaction productin the Ti/ZrO2 interfaces fabricated by hot pressing,whereas according to the thermodynamic analysis, it is pre-dictable that Ti can react with ZrO2 in the synthesis proc-esses of Ti/ZrO2 FGM, and the bonding state between Tiand ZrO2 is physical in the composite. The XRD results weredemonstrated that the volume fraction of Ti in the compo-sites has remarkable effect on the phase transformation

from t-ZrO2 to monoclinic zirconia (m-ZrO2). Under thesame sintering conditions, the volume fraction of m-ZrO2

increases with the increase of Ti content.114 In high-temper-ature sintering at 1500�C, studies with 90 mol % Ti showedthat ZrO2 particles were almost completely dissolvedin Ti, being accompanied by simultaneous precipitationof Y2Ti2O7.

115

Functionally graded ceramic–ceramic materialsGuo et al.88 described the successful preparation of HA/ZrO2 FGM which is useful for dental implants at 1200�Cwithin 5 min by the SPS method and found that equiaxial Y-TZP grains uniformly dispersed in the HA matrix. Theyshowed88 that there was no phase change of HA in the com-posites sintered at 1100�C, but HA started to decomposeinto a-TCP at 1200�C. On the other hand, the Y-TZPremained the t-ZrO2 phase even after the maximum temper-ature of 1250�C. A thin layer of CaZrO2 phase was found onthe interfaces among HA grains and zirconia grains. TheYoung’s modulus of these FGM grew with increasing zirco-nia content and sintering temperature up to 140 GPa sin-tered at 1100�C and 160 GPa at 1200�C, which is high com-pared to Ti/HA. Guo et al.88 reported that the bendingstrength of the composites SPS at 1200�C reached 200 MPa,which is double the strength of pure HA ceramics. It hasenough durable strength in the clinical practice. Matsunoet al.87 have also investigated HA/ZrO2 FGM by hot pressing,but the significant difference between the Guo study andthe Matsuno’s study was the design and arrangement of theindividual layers. Guo et al.88 designed the pure HA layer asthe middle layer and the HA þ 40% Y-TZP layers on bothends of a prepared disk as shown in Figure 6. This design’suse introduces compressive stresses that can be developedon both ends of the disks, and there were no long micro-cracks observed in their spark plasma-sintered samples. Themicrocracks are the direct evidence of thermal stressesowing to a cooling process from a high temperature toroom temperature and are generally present in both compo-sites and coatings. Guo et al.88 reported that the tensilestress developed in the HA layer because of its higher

FIGURE 6. Schematic diagram showing a cross-section of a disk

specimen of the HA/Y-TZP functionally graded composite, consisting

of symmetric layers with graded compositions.88


thermal expansion coefficient than that of the ZrO2 layer,and therefore causing the microcracking in the HA layer.

Porous FGM dental implantsThe simplest fabrication technique for manufacturing po-rous FGM dental implants is based on the partial densifica-tion during sintering of metal powders. This method isknown as powder metallurgy. Thieme et al.116 produced po-rous Ti FGM destined for orthopedic implants with a poros-ity gradient perpendicular to the surface by a powder met-allurgy technique. The porous FGM demonstrated adequateYoung’s modulus in the range of 5–80 GPa so that it isadapted to the elastic properties of bone with the purposeof avoiding stress shielding effects and to present betterlong-term performance of the implant-bone system. Theresults indicated that the modulus is inversely connectedwith the porosity gradient.116

Kutty and Bhaduri 103 developed a graded porosity onthe surface while maintaining a dense core in a Ti sam-ple.116 In their research, the synthesis was carried out in asemi-industrial grade microwave cavity using a-SiC suscep-tor. The Ti samples were sintered at 1, 1.25, and 1.5 kW for30 min. The result was that a pore size in the range of 30–100 lm is ideal for fibroblast ingrowth.12 A mechanicalstudy showed the highest strength of about 400 MPa whenthe samples were sintered at 1.25 kW for 20 min and thisvalue is close to bulk CP Ti (345–550 MPa). In addition,FGM porosity and the graded nature of change from densecore to porous surface should result in a better stress trans-fer than a coated surface. Therefore, a one-step processingtechnique works for fabricating various implants and dentalimplants. A pulsed electric current sintering (PECS) processexists for the production of porous structure with a porositygradient in the micrometer size.106 PECS permits morerapid sintering at lower temperature than conventional sin-tering, primarily by means of a spark pulse current thatforms between the powder particles.117

Suk et al.104 used a pressureless PECS method for manu-facturing porous FGM, and the obtained specimen has agradual change in pore distribution. This experiment wascarried out without rapid and strong neck formation andwith no volume shrinkage. The major interest in this pro-cess is where to obtain good pore interconnectivity, easy po-rosity control, and short processing time, but low porosityand expansiveness.12,104 Figure 7 shows the graphite moldused in the PECS technique for the preparation of the po-rous material with a porosity gradient. Three K-type ther-mocouples at various positions, as shown in Figure 7 (posi-tions 1–3), checked the temperature difference along thelongitudinal direction within the graphite mold. This processconfirmed that a suitable temperature gradient was estab-lished for preparing the porous structure with a porositygradient of reasonable degree. Sintering was generally per-formed with a thermocouple inserted completely into thehole at position 2, the temperature of which was taken asthe significant sintering temperature.

Laser-Engineered Net Shaping (LENSTM) is a techniqueable to obtain gradient structures and can also produce net-

shaped implants with designed porosities and can extend toother metallic biomaterials as well.118,119 Krishna et al.105

used LENS to produce porous Ti implants with mechanicalproperties matching those of natural bone. They reportedthe bulk density and porosity of these samples varied,depending on the LENS-processing parameters.105 Young’smodulus increased linearly with increasing density of thesamples. The experimental data indicate that the modulusand strength of laser-processed Ti specimens can be tai-lored in the range of 2–45 GPa and 21–461 MPa, respec-tively, whereas the porosity was in the range of 35–42 vol%, close to those of human cortical bone. They alsoobserved that it is possible to fabricate porous Ti sampleswith a porosity higher than 40 vol % that have appreciablemechanical properties, whereas in other literatures, it wasreported that the strength of compacts with a porosityhigher than 40 vol % is diminished to approximately 0MPa.120,121 Generally, the important points in this processare good pore interconnectivity, control over porosity, andability to realize complex shapes. However, this technique islimited to small pore sizes.

More recently, the development of direct laser metal sin-tering (DLMS) processes has considerably increased thefield of application of Ti alloys and allowed implants to befabricated more economically in comparison with traditionaltechniques. Among the several direct metal-forming techni-ques, selective laser sintering promises great potential bene-fits in the field of the biomaterials, especially in implantdentistry, owing to its capability to directly build three-dimensional (3D) metallic components from metal powderwith minimal or no postprocessing required.122,123 Trainiet al.93 used DLMS and subsequently modified it with acid-etching methods to improve the surface microstructure forthe production of isoelastic FGM for porous Ti dentalimplants. In their research, the original surface consisted ofspherical particles in the range of 5–50 lm. In the nextsteps, they used different etching techniques using inorganicand organic acids. It was observed that after exposure tohydrofluoric acid, some of the spherical particles wereremoved and the microsphere diameter was in the range of5.1–26.8 lm. After an organic acid treatment, grooves from14.6 to 152.5 lm in width and 21.4–102.4 lm depth

FIGURE 7. Schematic illustration showing the graphite mold used in

the PECS process: (1) specimen (powder), (2) die, and (3) punch.

Holes for inserting thermocouples are located at positions 1–3.104

REVIEW ARTICLE


replaced particles. The SEM images of the surface’s DLMS-fabricated disks after different treatments are shown inFigure 8. Moreover, they assessed the variation of surfaceroughness by use of a Chi-square test. It is important tocontrol the surface morphology of the dental implants toobtain the desired surface roughness and the wettability toallow the absorption of plasma proteins on the implant sur-face.124 Young’s modulus of the exterior porous Ti was 776 3.5 GPa, but the mean of porosity was 28.7% of the metalsurface, and that of the inner core Ti was 104 6 7.7 GPa.The fracture face indicated a dimpled appearance typical ofductile fracture. However, the method is also limited in thearrangement and the control of pore size that can befabricated.

FUTURE CONSIDERATIONS

For nearly all dental implants, the main goals are rapidreturn to function (i.e., mastication), stronger, safer osseoin-tegration, and long-term fixation of implants to bone. Toachieve these goals, designers of dental implants must con-front biomechanical and biomaterials subproblems, includ-ing in vitro and in vivo performances on implants, mechani-cal compatibility to smooth transfer the stress between theplaced implant root and the receiving hard tissue, and inter-facial tissue response. As a promising candidature, FGM canbe applied to improve the success of dental implants whilethere are no many examples of actual FGM dental implantsthat have been used in clinical applications. The hurdle maybe the producing cost and transferring the techniques toshape of a dental implant, which will prevent the realizationof the potential of functionally graded dental implants. Inour research to create the ideal implant designed withFGMs approaches, we are required to close key gaps in ourbasic knowledge and make a series of prototype dentalimplants with increasing functionality. We must start by dis-tinguishing the specific gaps, in our present knowledge, andthen look widely for advances that will facilitate us to linkthem. We must try to uncover the relationships linking com-position and biomaterials architecture in functionally gradeddental implants with mechanical behavior, length scales, andthe capability for osteogenesis. Finally, these relationshipsshould be tested and evaluated systematically in vivo andfinally in clinical studies.

Although many biomaterials and ceramics can be madeinto the FGM and gradient porous dental implants, but sofar they have received little attention and still need forimprovement. For example, recently bioactive glass scaffoldswith strengths comparable to those of cortical bone havebeen produced. Therefore, these bioceramics may havepotential for the FGM in dental implants. But, the mechani-cal properties and reliability of BG remain as limiting fac-tors for applications in dental implants; however, still moreconsideration is needed for development.

The study of the literature suggests that many differentparameters might be important for the long-term osseointe-gration and better performance in bone, where the goal canbe achieved by lowering the FGM material gradient. But, itis well established that this will, at the same time, decreasethe stiffness of implantation, as a result locating the boneimplant interface at higher risk of damage during the earlyhealing stage. The problem might be minimized by the mul-tiobjective optimization processes that are yet to be used inbiomechanical studies up to date, whereas machining, den-tal, and material demands must be considered. There is aneed to further analyze and further information on themachining characteristics and machining-induced surface ofdental bioceramics and FGM dental implants to improvequality, efficiency, bone–implant contact, and costproduction.

Depending on the clinical advantage required for,merged with the regulatory landscape, the commercial de-velopmental efforts will finally provide new products withnovel and unique properties.

CONCLUSIONS

Close studies of natural tissue and organs illustrate thatthey are not homogenous and natural functional gradientsexist in their formation. Biomaterial and tissue engineeringresearch literature shows that there is a requirement indeveloping implants with FGM. Therefore, the great poten-tial of FGM lies not only in the field of bone tissue engineer-ing but also in dentistry. To improve the acceptance of den-tal implants inside the living bone, the concept of FGM is afavorable approach. The works reviewed here show that thedevelopment of functionally graded dental implants and po-rous functionally graded dental implants is feasible for

FIGURE 8. SEM images at 500� magnification of surface’s fabricated disks after different treatments. (a) Untreated disk; (b) disk treated with

hydrofluoric acid; and (c) disk treated with organic acid.93


commercial dental implants. Porous functionally gradeddental implants were reviewed to improve mechanical prop-erties, especially Young’s modulus as the high stiffness of Tigenerally is not well-matched to bone, which can lead tostress shielding of the residual bone, which in turn mayresult in detrimental resorptive bone remodeling.

There are a variety of techniques for producing FGMdental implants and gradients of porosity. However, thesestructures are largely still in the stage of laboratoryresearch. The obstacle may be the producing cost, whichwill prevent the realization of the potential of FGM dentalimplants. Also, more tests and studies are needed withregard to design and production of combinations of bioinertand bioactive materials for FGM for dental implants.

ACKNOWLEDGMENTS

The authors are grateful for the grants.

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Mehrali Et Al. [2013] Dental Implants From Functionally Graded Materials

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