Nanotecnologie per il miglioramento degli impianti dentali

Nanotechnology Approaches for Better Dental Implants

Antoni P. Tomsia1, Maximilien E. Launey1, Janice S. Lee2, Mahesh H. Mankani3, Ulrike G.K.Wegst4, and Eduardo Saiz5

1 Materials Sciences Division, Lawrence Berkeley National Laboratory, Berkeley, CA 947202 Department of Oral Maxillofacial Surgery, University of California, San Francisco, 513Parnassus Avenue, San Francisco, CA 941433 Department of Surgery, University of California, San Francisco, 1001 Potrero Avenue, Box0807, San Francisco, CA 941434 Department of Materials Science and Engineering, Drexel University, 3141 Chestnut Street,Philadelphia, PA 191045 Centre for Advanced Structural Ceramics, Department of Materials, Imperial College London,Exhibition Road, London, UK

AbstractThe combined requirements imposed by the enormous scale and overall complexity of designingnew implants or complete organ regeneration are well beyond the reach of present technology inmany dimensions, including nanoscale, as we do not yet have the basic knowledge required toachieve these goals. The need for a synthetic implant to address multiple physical and biologicalfactors imposes tremendous constraints on the choice of suitable materials. There is a strong beliefthat nanoscale materials will produce a new generation of implant materials with high efficiency,low cost, and high volume. The nanoscale in materials processing is truly a new frontier. Metallicdental implants have been successfully used for decades but they have serious shortcomingsrelated to their osseointegration and the fact that their mechanical properties do not match those ofbone. This paper reviews recent advances in the fabrication of novel coatings and nanopatterningof dental implants. It also provides a general summary of the state of the art in dental implantscience and describes possible advantages of nanotechnology for further improvements. Theultimate goal is to produce materials and therapies that will bring state-of-the-art technology to thebedside and improve quality of life and current standards of care.

1. IntroductionReplacement of tooth and bone with metal implants and plates is one of the most frequentlyused and successful surgical procedures. The introduction of modern implants started withthe work of Branemark, who in 1969 observed that a piece of titanium embedded in rabbitbone became firmly attached and difficult to remove [1]. This led to Branemark’s carefulcharacterization of interfacial bone formation at titanium implant surfaces, anddemonstration of excellent osseointegration [2, 3]. Dental implants based on the Branemarkwork were introduced in 1971 [4]. Thetotal hip replacement was first developed by Charnleyin 1962, with stainless steel used for the stem and the attached ball [5, 6]. Charnley’s maincontribution was to adapt PMMA cement from dentistry to fix the femoral stem andacetabular cup [7]. After decades of subsequent research in industry and academia, implants

Corresponding author: Antoni P. Tomsia, PhD, Materials Sciences Division, Lawrence Berkeley National Lab, One Cyclotron Road,MSD Building 62R0203, Berkeley, CA 94720, Phone: 510-486-4918, [email protected].

NIH Public AccessAuthor ManuscriptInt J Oral Maxillofac Implants. Author manuscript; available in PMC 2011 May 5.

Published in final edited form as:Int J Oral Maxillofac Implants. 2011 ; 26(Suppl): 25–49.

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have evolved, with a high percentage of survival rate and longevity. The research resulted inbetter design, better materials, and more extensive clinical experience compared with theearly years of implant development. One can argue, however, that the progress has beenslow and incremental. We are still using almost the same alloys described in the works ofBranemark and Charnley, many of which were originally developed by the aerospaceindustry. Scientific evidence to determine the success (or failure) of dental implants in theUnited States is almost nonexistent, as we lack a national registry describing the rate andcauses of failures. There are, of course, other avenues to alert clinicians about implantfailures: the peer review literature, other national registries (the American Academy ofOrthopaedic Surgeons (AAOS) is sponsoring one in the United States for hip and kneeimplants), Food and Drug Administration alerts, manufacturing recalls, implant-retrievallabs across the country and world, implant-retrieval studies in collaboration with clinicalfollow-up investigations, class-action lawsuits, and society presentations. The developmentof retrievable information relevant to medical devices is necessary to understand the overallperformance of a particular implant and to improve the next generation of medical implants.For implant-design improvements, implant-retrieval programs should be coordinated withclinical-outcome studies that include a patient’s history and medical records. Implant-failureanalysis not only helps manufacturers design better products but also provides betterunderstanding for clinicians on how the implant interacts with the body and which ones arebest for a patient’s particular situation. In 2003, more than 700,000 dental-implantprocedures were performed in the United States and more than 1.3 million procedures inEurope [8]. Approximately1 million artificial hips and knees are implanted each year in theUnited States, as reported in 2006 [9]. Hip and knee replacements have a success rate ofmore than 90%, and dental implants fare better, at 90–95% [10].

Dental implants have a long and successful history. The percentage of failure is very low,approximately 5%, mostly likely due to infection, rejection, accelerated bone loss, and poorosseointegration with loosening of the implant [10]. The most frequent cause for failure isinsufficient bone formation around the biomaterial immediately after implantation [11].Therefore, this is an area where improvements are needed. This is becoming even moreimportant as clinicians are pushing for faster healing times. Consequently mostmodifications in the implant design are geared toward reducing the time needed to waitbefore loading. It is here that the implant surface and tissue interface are critical [12].Current surface chemistries and morphologies are controlled, at best, at the micron level, buttissue response is mainly dictated by processes controlled at the nanoscale. Understandingand controlling interfacial reactions at the nano level is the key to developing new implantsurfaces that will eliminate rejection and promote adhesion and integration to thesurrounding tissue.

Here, the first promising avenue is through surface engineering that would include nanoscaletopography and/or coatings for better and faster osseointegration of implants designed tofunction for the lifetime of the individual. However, the literature describing and comparingthe different approaches is often confusing. Worldwide, there are 80 companies producing220 various dental implants, so the comparison is often difficult [10]. Likewise, there arefew controlled studies simultaneously comparing different implants, especially in theclinical setting. Manufacturers often don’t find it in their best interest to participate in head-to-head competitions with their rivals, and clinicians often don’t have sufficient patientnumbers or motivation to initiate these studies on their own. Additionally, there arevariations in the patient population that may influence the success of the implant. There areno commonly accepted guidelines for the design of surface topographies or the deposition ofcoatings, and the clinical evidence is often conflicting. As described later, this may be due tothe poor control of factors associated with fabrication techniques and not necessarily withthe design itself. One critical problem is how to translate the results of in vitro cell-culture

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studies performed on model surfaces in pristine conditions to the design and fabrication ofthree dimensional (3D) implant materials with complex shapes that will work in the harshhuman oral environment.

The second scientific challenge is to develop a new generation of implant materials that willcombine the advantages of ceramics, in particular their inertness, with a mechanicalresponse comparable to those of dental implant alloys. The unique properties of ceramicmaterials, including their outstanding corrosion resistance and excellent aesthetics, makethem appealing candidates for many dental applications. However, their inferior mechanicalproperties, particularly fracture toughness, until now have hampered their commercial use,especially in load-bearing situations. This is a particular problem when designing porousimplants or implants with rough and porous surfaces for better osseointegration. While theporosity will improve osseointegration, increased porosity can easily result in diminishedstrength and fracture at relatively low loads. Through new fabrication technologies, theselimitations may be overcome. Their potential has not yet been realized.

Solutions to these scientific challenges have a benefit that extends beyond the constructionof implants. They can offer us a path toward the resolution of a much more challenging goal:the construction of a true biologic tooth. The technical foundation for a neo-tooth will likelyinclude implant surfaces that bind quickly to bone, and occlusive surfaces that offertoughness during mastication. In the meantime, we need strategies to improve the currentmetallic dental implants, through surface modifications of the implant either by applyingnovel ceramic coatings or by patterning the implant’s surfaces. These approaches aredescribed below.

2. Surface modifications of dental implants2.1 Ceramic coatings

Various coatings have been developed to improve an implant’s ability to bond to livingtissues, particularly bone. The idea is to apply a thin ceramic layer that will bond both to theimplant and to the surrounding tissue while promoting bone apposition. Candidate coatingmaterials are bioactive compounds able to promote cell attachment, differentiation, and boneformation. The most prevalent bioactive materials are calcium phosphates (CP), such ashydroxyapatite (HA) or tricalcium phosphate (TCP), and bioactive glasses. When implanted,these bioactive ceramics form a carbonated apatite (HCA) layer on their surfaces throughdissolution and precipitation. This phase is equivalent in composition and structure to themineral phase of osseous tissue. At the same time, collagen fibrils can be incorporated intothe apatite agglomerates. The sequence of events is poorly understood but appears to be:adsorption of biological moieties and action of macrophages, attachment of stem cells anddifferentiation, formation of matrix, and, finally, complete mineralization [13–17].

Animal studies have shown that, compared with uncoated implants, implants with a thinlayer of CP have markedly enhanced interfacial attachment of bone tissue over a period ofseveral months [19–26]. Also, numerous histological studies provide evidence that coatedimplants yield a more reliable interface with bone than mechanical osseointegration of Ti[27–36]. Coating implant alloys with CPs, therefore, is advantageous; it facilitates joiningbetween a prosthesis and osseous tissue and consequently increases long-term integrity.Benefits arise during healing and subsequent bone-remodeling processes [30, 37], includingfaster healing time [38], enhanced bone formation [39, 40], firmer implant-bone attachment[39, 41], and a reduction of metallic ion release [42, 43]. However, coating dental implantsis not without controversy. Recent clinical orthopedic studies [44–46] indicate that there isno clear advantage to using HA coatings for the femoral component of hip implants. Thesituation with dental implants is no different; there are contradicting reports on the



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effectiveness of the HA coatings, as it is difficult to compare coatings applied using differentmethods [47]. Thus, the issue of coating usefulness is not fully resolved.

Ideally, the coatings should be tailored to exhibit prescribed biological attributes for eachspecific application. They should have strong adhesion to the implant and good fixation tobone. Their microstructure and dissolution rates must be programmed to match the in vivohealing process, and they could serve as templates for the in situ delivery of drugs andgrowth factors at the required times. Essential requirements of the new family of coatingsinclude: (1) good adherence to the implant, (2) fixation to the bone, (3) fine gradient andthickness control with programmed dissolution rate in body fluids, and (4) therapeuticcapabilities. These are discussed below.

Coating/implant adherence—Most reports indicate that currently available coatingtechniques provide inadequate adherence of CP coatings to Ti alloys. Quoted bondstrengths, 15–30 MPa, are very low [48–51]. Systematic studies of the chemical andmicrostructural factors that control the interfacial strength and toughness are needed tooptimize the mechanical stability of the coatings. Key principles for bonding ceramics tometals, elucidated over the past decade [52–63], can guide such an investigation but we stillneed to develop standard procedures to test adhesion, and a clear understanding of themechanisms of interfacial failure in the body environment.

Fixation to bone—Knowledge of biological processes occurring at the biomaterial-tissueinterface is of utmost importance in predicting implant integration and host response [64].Two primary modes of attachment are: (1) mechanical, in which the implant has a rough,porous surface into which bone grows (often supplemented by use of a bone cement); and(2) chemical, in which bone “bonds” to the implant material. Here it is difficult to decouplethe effect of coating chemistry and topography. It has been observed that differences in thenature of the implant/bone contact depend upon the specific coating. However, there is littleguidance as to which is the ideal coating chemistry. Dissolution of calcium and phosphorusfrom the coating may promote mineralization and bone formation but it is not clear howmuch coating solubility contributes to the best fixation, and excessive resorption can limitimplant lifetimes [65, 66].

Programmed dissolution rates—A critical goal in the design of novel coatings is theprogramming of their dissolution (bioresorption) rates. The role of solubility of coatings inthe body is poorly understood [67]. The presence of highly soluble phases markedlydecreases the mechanical stability of the coating in vivo, but some solubility of coatingmaterial expedites fixation [68, 69]. These results suggest that graded coatings designedwith a soluble surface to facilitate bonding to bone and an insoluble layer in contact with themetal to provide adhesion, corrosion resistance, and long-term mechanical stability couldoffer a significant improvement over current materials. Both the composition and thicknessof the graded coating layers can be controlled to manipulate the resorption rates. Theirresorption can be programmed to match healing rates and to expose different micro-architectures, chemical patterns, and porosities at different times to optimize the biomaterialcoating surface for different periods of the healing-in phase. Possible designs of gradedcoatings are depicted in Figure 2.

Drug delivery—Inflammatory responses to implants are a significant problem in the healthcare industry. Typically, medications are given to a patient following surgery to suppressinflammation and to enable the intended performance of the implanted medical device.Although generally helpful, in many cases this approach is insufficient or entirely ineffective[70–72]. A different approach is to provide a local dose of anti-inflammation agentsgradually released from a coating on the surface of the implanted device [73, 74]. The main



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advantage of this approach over traditional means of administering the drug is that the drugcan be directly released at the implant site without having to go through the bloodstream.This lowers the amount of drug needed, reducing the overall toxicity and side effects. Inaddition, growth factors that are known to encourage tissue-implant integration, such asTGF-β [75, 76], may be delivered locally using this platform. These chemicals could beincorporated into the porosity of micro- and nanoporous coatings or could be dispersed inbiodegradable polymers, and combined with the bioactive glass/CP coatings. Developing theright platform to control the release rate is the key, and closely related to the control of thedegradation rates.

Perhaps the most popular coating method is plasma spraying of HA [77–82]. However,plasma spraying of HA offers only rudimentary control of the coating thickness andcomposition. The thickness of plasma-sprayed coatings typically exceeds 50 μm. The highprocess temperatures cause partial thermal decomposition of HA, leading to the formation ofother CPs, including 22–62% of highly soluble amorphous calcium phosphate [83–87], α-TCP, β-TCP, tetracalcium phosphate, and calcium oxide [49, 88–90]. The result is coatingswith unacceptable heterogeneous properties. Long-term in vivo studies demonstrate thatsome of the major causes of failure of HA-coated metal orthopedic and dental implantsreside in the coating [47, 91]. Severe problems include: (1) unreliable adhesion—they candetach, yielding floating fragments [92–99]; (2) continued dissolution of the coating, leadingto catastrophic failure of the implant at the coating-substrate interface [100]—attempts toincrease crystallinity and reduce solubility by post-coating thermal annealing markedlydegrades the coating adherence [88, 91, 101–104]; (3) variability during processing in termsof resulting phases, stresses, and cracking [37, 84, 105–108]; (4) partial crystallization of thedeposited coating, leading to the presence of more soluble CP compounds [48, 109–111]; (5)significant degradation in the fatigue resistance and endurance strength of the implant alloy[112]; (6) very irregular morphology of the coatings; (7) poor control of the coatingthickness, with greater risk of fracture, the thicker the coating; and (8) poor control ofphysicochemical properties, and hence the biological stability of the coating [113, 114].

Several other techniques have been used to apply CP coatings on metallic alloys, from sol-gel to RF magnetron sputtering and others [106, 114–132]. These techniques may offer amore accurate compositional control and the possibility of fabricating much thinner layers(of the order of 1 micron or less). This could be advantageous for coating stability, as thedriving force for cracking and delamination decreases with the decreasing thickness. Thenature of the coating and the process might be tuned in order to modify the interfacialstrength. However, these approaches have yet to be translated into commercial designs.Some are line-of-sight techniques not suitable to coat implants with complex shapes,resulting in incomplete coverage of the implant; cost is a key variable that must be takeninto account and highly dense coatings produced by sputtering are probably the mostdesirable, but are very expensive to fabricate.

There is increasing interest in the use of bioactive glasses in the fabrication of coatings. Thepotential of bioactive glass coatings has been recognized since the discovery of the originalBioglass ® composition by Hench [90]. Bioglass® has excellent bioactivity and could beused to enhance the adhesion of the implant to the bone. Furthermore, the glass propertiescan be easily adjusted from bioactive to bioresorbable or bioinert by controlling thecomposition [90]. This creates the possibility of fabricating layers with finely tunedproperties. However, initial attempts to coat metallic implants with the original Bioglass®

were marred by the generation of large thermal expansion stresses and high reactivitybetween the metal and the glass [16]. Subsequently, several techniques have beeninvestigated for the preparation of glass coatings, such as enameling, plasma spray, RFsputtering, sol-gel, pulsed-laser deposition (PLD, Figure 3), and others [47, 133–141].



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Although it has been possible to deposit glass layers that retain bioactivity, a recurrentproblem is the control of the glass-metal reactions to achieve good coating adhesion, and thegeneration of large thermal stresses during processing that can cause cracking and/ordelamination of the coating. The solution to the problem requires the development ofcoating procedures and novel bioactive glasses, and glass-ceramics, with adequate thermalexpansion coefficients and softening points.

In our laboratory, we have demonstrated that a new family of bioactive glasses, in thesystem Si-Ca-Mg-Na-K-P-O, exhibits thermal-expansion coefficients close to those of themetallic alloys used in dental and orthopedic implants [142]. This is achieved through thepartial substitution of Ca by Mg, and of Na by K. These glasses have adequate softeningpoints to be used in the fabrication of coatings by a simple enameling technique. Strongglass/metal adhesion is achieved through the formation of thin interfacial layers. Theprocedure can be extended to the fabrication of graded coatings through the sequentialdeposition of different glass and glass-ceramic layers. For example, we have fabricatedgraded coatings on Ti and Co-Cr alloys consisting of a high-silica glass in contact with themetal, and a composite surface layer that consists of a mixture of a low-silica glass withsynthetic HA particles. The high-silica layer is very resistant to corrosion in body fluids; itprovides good coating adhesion and long-term stability, while the surface layer is designedto enhance coating attachment to the tissue (Figure 4).

The interaction of glass coatings with tissue can be further manipulated by designingappropriate surface textures. Furthermore, in vitro analysis in cell cultures has shown thatbioactive glasses and coatings can indirectly affect osteoblast gene expression through theirdissolution products [143–147]. Quantitative real-time reverse transcription-polymerasechain reaction (RT-PCR) analysis showed that the silicate glass coating extract induced atwofold expression of Runx-2, a key marker of osteoblast differentiation, compared withTi-6Al-4V and tissue-culture polystyrenes [148]. Future research should address theidentification of specific chemical cues behind the induction of gene expression, and thetailoring of layer compositions to control such effects.

Another technique worth further exploration is anodic oxidation, which has been used as ameans to engineer the surface of Ti-based implants [149, 150]. This technique can be used tocreate an adherent oxide coating on the implant surface. By selecting different electrolytesand manipulating the conditions, it is possible to create oxide layers with a wide range ofstoichiometries as well as micro- and nanoporosities. It is also possible to incorporate Caand P ions into the layers. This is a relatively simple and economical technique that can beeasily adapted to fabrication, and anodized implants are commercially available. Several invivo and clinical studies have shown that anodization enhances integration at early stagesand due to their characteristic micro to nano pore distribution, anodized surfaces have beenconsidered as platforms for drug delivery. As in many other treatments, it is difficult here toseparate the effect of surface chemistry and morphology. In addition, as in other implanttechnologies, we still lack systematic studies of the factors that control adhesion of the oxidelayer to the alloy.

Most likely an ideal coating will not be fabricated using a single procedure or material, butby combining several of them to fabricate layers that blend organic and inorganic phases,with thicknesses ranging from hundreds of microns to the nanometer level, and chemicallyand topographically textured surfaces. Each technique has specific advantages andlimitations, yet no single procedure can fabricate layers combining organic and inorganicmaterials and generate textured surfaces. New coating techniques should be flexible enoughto use glasses and CP phases with a wide range of compositions and should generatechemically and topographically textured surfaces while integrating organic and inorganic



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components. To accomplish this, colloidal techniques (dip-coating, electrophoreticdeposition, EPD, or robocasting) able to prepare thick layers (>10 μm) can be combinedwith thin-film formation using techniques such as pulsed laser deposition (PLD) (Figure 3)[151, 152]. The translation of rapid prototyping techniques to the fabrication of coatingscould prove very useful. For example, robocasting, a 3D technique that can print differentmaterials at the same time and produce porous surfaces for infiltration of organic materials,can also be employed. These techniques can also be used to integrate organic materials intocoatings to serve as templates for the delivery of drugs and growth factors.

2.2 Surface functionalizationAn alternative or complementary strategy to the use of coatings for dental implants is themolecular grafting or chemical treatment of implant or coating surfaces to enhance celladhesion and promote mineralization, and the production of matrix and marker proteins.Different treatments have been used to create hydrophilic implant surfaces (e.g., throughchemical etching) to promote protein adhesion and subsequent cell attachment (Figure 5).However, the native oxide layer usually present on the surface of metallic implants exhibitsrelatively low contact angles with water without any particular treatment. Overall, several invivo studies suggest that implants with increased hydrophilicity exhibit faster integrationwith larger bone-to-implant contact, but key questions remain: Is hydrophilicity the keyvariable, and if so what degree of hydrophilicity is needed to make a difference? Are thereadditional benefits derived from the surface treatment?

The logical candidates for molecular grafting are proteins present in the extracellular matrix,and implant surfaces that have been functionalized with fibronectin (FN), vitronectin (VN),or laminin (LN), to name a few [153]. Lately, the focus has shifted to the use of signalingdomains, composed of several amino acids. These domains are present along the chain ofthe extracellular matrix proteins and are the ones that interact with cell-membrane receptors.Perhaps the best-known example is Arg-Gly-Asp (RGD), the signaling domain derived fromFN and LN. However, other sequences such as Tyr-Ile-Gly-Ser-Arg (YIGSR) or Arg-Glu-Asp-Val (REDV) have also been used [154–156]. The use of a short peptide is moreconvenient because the long molecules can be folded, and as a result the binding domainsmay not be available. Two important aspects are the development of techniques to bind themolecules to the surface and the control of their spatial distribution.

Several approaches have been used to functionalize Ti surface with different molecules,from adsorption (physical or chemical) to the use of covalent bonding or self-organizedlayers [157–159]. The simplest procedures involve immersion of the material in a solutioncontaining the desired molecules in order to promote adsorption on the surface. Thisapproach has disadvantages because it can result in relatively poor adhesion. Covalentbonding using, for example, linker molecules can promote stronger adhesion to the implantand it has been used to attach proteins, signalling domains [157–159], antibiotics [160], andgrowth factors such as human epidermal growth factor or recombinant human bonemorphogenetic protein-2 to Ti and TiO2 surfaces [161, 162]. However, it may impose aspatial orientation on the grafted domains or proteins, limiting their effectiveness [154].Diverse molecules can also be incorporated into CP coatings or on Ti surfaces prepared byanodic polarization [157–159]. The surface density is also important. It has been shown thata way to enhance the function of the osteoblastic cell lines is to match the surface density ofthe selected molecules to the distribution of the corresponding receptor in the cellmembrane. This can be achieved, for example, using functionalized gold nanoparticleswhose density on the surface can be manipulated [163, 164]. Although most of the paperssupport the idea that functionalization can enhance and accelerate osseointegration, the keyis to adapt the functionalization techniques to the surfaces of specific dental alloys, each onewith different surface chemistries, and to be able to create surfaces that will display different



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molecules with controlled spatial distribution and the required degree of adhesion.Furthermore, we probably still lack the in vivo data needed to select optimum attachmentmethods and surface densities.

2.3 Surface topography of dental implantsIt has long been recognized that the chemical and morphological characteristics of implantsurfaces affect the behavior of the surrounding tissues. Numerous studies have shownmarked differences in the in vitroand in vivoresponse of textured implant surfaces,demonstrating that the ability of the implants to support bone formation can be enhanced bymodifying surface topography [165–168]. This is critically important in the case of dentalimplants that are built using bioinert materials and where optimum and fast osseointegrationis needed to achieve long-term functionality. It is generally accepted that rough surfaces willfavor mechanical interlocking with the surrounding tissue, which will help implant stability.At a more fundamental level, surface topography may affect cell differentiation (perhaps byinfluencing cell shapes or by inducing specific signals) [169–171] and consequently promotethe formation of extracellular matrix and bone apposition. Although there are hundreds ofpapers dealing with surface engineering, decoupling and understanding these differenteffects is still an open scientific challenge. However, the answer could prove extremelyuseful in order to rationally design implant surfaces.

The topography of the implant surfaces can now be manipulated at a wide range of lengthscales, down to the nano level. Several implant designs use surfaces with large pores topromote bone ingrowth and favorable anchoring. Previous analysis indicated that thesepores have to be of the order of 100 μm and larger to allow for bone ingrowth. These can befabricated using partial sintering of metal particles or spheres on the surface. However, thereis still some controversy regarding the benefit of this approach with respect to screwgeometries. A recent study supports the idea that screw implants perform better in longlengths and denser bone while the porous surfaces are more adequate at short lengths and incancellous bone [172]. In addition, excessive roughness can have the disadvantage ofmaking the implant harder to remove if necessary, and may be detrimental for implantstrength [172].

Texturing of the implant surface at the micro level has been proposed as a feasiblealternative to promote bone apposition. This work is based on the abundance of in vitrostudies indicating that some degree of roughening at the microscopic scale favors cellattachment and differentiation [173, 174]. There are several papers and patents describingdifferent techniques to increase the surface roughness, including sandblasting, grinding, Tiplasma spraying, or laser texturing [28, 149, 175]. The in vivo results are mixed. Whilemany reports indicate an improvement in bone apposition using rougher surfaces, otherssuggest either no significant benefit or that the result will depend on the implant design. Forexample, Jansen compared the tissue reaction to smooth and microgrooved implants (width2 or 10 μm, depth 1 μm), at different implantation sites [176]. They were unable to prove theexistence of an effect of implant-surface topography on in vivo soft-tissue response. Otherrecent studies reached similar conclusions [177]. It could be argued that the main advantageof microtextured surfaces is the contribution to mechanical interlocking and tissueattachment of the implant. It should also be noted that implants with surfaces too rough havea high possibility of peri-implantitis and that the roughness will affect the degree ofhydrophilicity. This effect could be described using Wentzel equation cosθr = r cosθ (whereθr is the macroscopic contact angle on the rough surface, θ is the true contact angle, and r isa ratio between actual and apparent area, r ≥ 1) [178]. Notice that according to this equation,for θ < 90° as can be expected for Ti-alloy surfaces, roughness tends to decrease the contactangle (increase hydrophilicity). This simple model shows how roughness can be playing aneffect at multiple levels [179].



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Unfortunately, there are not many studies systematically comparing the performance ofimplants with surfaces treated using various methods. In particular, similar roughness(average peak-to-valley distances) can be obtained using diverse technologies but thecharacteristic of the surface (random vs. oriented features, sharp vs. smooth) can greatlyvary from one to the other and there is no standard approach to fully quantify the implanttopography and compare with in vitro and in vivo results. Furthermore, the surfacechemistry can also be altered in different ways by the roughening treatments. Ti-alloys arecovered by a native oxide layer whose characteristics can change with the treatment. Forexample, some studies suggest that sandblasting with TiO2 particles elicits more beneficialresults than using other ceramics [149, 180]. This is perhaps due to the presence of ceramicparticles left embedded in the metal after the sandblasting process. Finally, differenttreatments often result in a hierarchical surface structure where nanoscale features are alsopresent. These features are often ill characterized but, as we will see below, can have a largeimpact on the biological response.

The arrival of nanotechnology has opened new opportunities for the manipulation of implantsurfaces. It is believed that implant surfaces could be improved by mimicking the surfacetopography formed by the extracellular matrix (ECM) components of natural tissue. TheseECM components are of nanometer scale with typical dimensions of 10–100 nm. Many invitro studies have shown how cell attachment, proliferation, and differentiation areresponsive to nano-scale features such as pillars or grooves prepared, for example, usingnanolithography. In this respect, not only the size of the feature but also its distribution(ordered vs. random) can play a role. Nanopatterned surfaces may also provide betteradhesion of the fibrin clot that forms right after implantation, facilitating the migration ofosteogenic cells to the material surface [181]. At a more basic level, it is still not completelyclear that nanopatterning will be substantially better than patterning at the micron scale orwhat the interplay between surface topography and chemistry is. It is well known that a highdensity of nanopillars will create a superhydrophobic surface that can be detrimental [182]and many of the basic studies have only been performed on flat surfaces and using modelmaterials (e.g., silicon or PMMA) [183, 184]. How to extrapolate these techniques to thefabrication of dental implants, how to ensure that these features will be robust enough tosurvive implantation, or how to reliably test in vivo effects of nanoscale manipulation arestill open technological challenges.

3. Nanocomposites for bone regenerationDue to their strength and toughness, metal implants have been used in orthopedic and dentalsurgeries for many years. Ti and its alloys have had considerable advantages over othermetals because of their inertness, which yields excellent biocompatibility andnonsensitization of tissues. However, issues concerning the release of Ti and alloyingelements from implants and the formation of Ti debris due to wear during implantation stillremain. Metal and metal alloys (i.e. Ti, Al, V, Ni) in implants and dental bridges have thepotential for allergic reactions and vague constitutional symptoms [185, 186]. Ceramicmaterials are known to have excellent aesthetics, corrosion resistance, and biocompatibility;several ceramic implants have been already commercialized. The continuing interest in theuse of modern ceramics for the fabrication of dental implants is underscored by severalpresentations during recent meetings of the International Association of Dental Research,and by recent research efforts by several European and Japanese companies (Kyocera,Dentsply, Metoxit, et al.). Unfortunately, in contrast to metallic materials, most ceramicssuffer from almost a complete lack of plastic deformation; this is due to the absence ofmobile dislocation activity, although other modes of inelastic deformation, such asmicrocracking and in situ phase transformation, can provide limited alternative deformationmechanisms. The implications from this are that ceramics are inherently brittle, with an



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extreme sensitivity to flaws. As a result, fracture almost invariably occurs catastrophically(with crack initiation concomitant with instability) by cohesive bond breaking at the cracktip with a resulting very low (intrinsic) toughness of roughly 1 to 3 MPa√m. This problem isaggravated when designing porous implants or implants with rough and porous surfaces forbetter osseointegration. In this case, the porosity can easily result in diminished fracturestrength at relatively low loads.

This dependence of strength on pre-existing flaw distributions has several importantimplications for brittle materials. In particular, large specimens tend to have lower strengthsthan smaller ones, and specimens tested in tension tend to have lower strengths thanidentically sized specimens tested in bending because the volume (and surface area) ofmaterial subjected to peak stresses is much larger; in both cases, the lower fracture strengthis associated with a higher probability of finding a larger flaw. Toughening ceramics, aswith virtually all brittle materials, must be achieved extrinsically, i.e., through the use ofmicrostructures that can promote crack-tip shielding mechanisms such as crack deflection,in situ phase transformations, constrained microcracking (although this mechanism isgenerally not too potent), and, most importantly, crack bridging. Extrinsic mechanismsresult in resistance-curve (R-curve) behavior (where the crack-driving force to sustaincracking increases with crack extension) as they operate primarily behind the crack tip tolessen the effective crack-driving force; they are therefore mechanisms of crack-growthtoughening. The incorporation of reinforcements in the form of fibers, whiskers, or particlescan also toughen ceramics, although the motivation may be rather to increase strength and/orstiffness. For toughening, crack bridging is the most prominent mechanism, particularly inceramic-matrix composites; by utilizing fibers with weak fiber/matrix bonding, when thematrix fails, the fibers are left intact, spanning the crack wake and acting as bridges toinhibit crack opening [187].

The key question is: Can we create a new family of ceramic-based materials that will matchor improve the mechanical performance, in particular the combination of strength andtoughness, of metallic alloys? In ductile materials such as metals and polymers, strength is ameasure of the resistance to permanent (plastic) deformation. It is defined, invariably inuniaxial tension, compression, or bending, either at first yield (yield strength) or atmaximum load (ultimate strength). The general rule with metals and alloys is that thetoughness is inversely proportional to the strength. In brittle materials such as ceramics,where at low homologous temperatures macroscopic plastic deformation is essentiallyabsent, the strength measured in uniaxial tension or bending is governed by when the samplefractures. Strength, however, does not necessarily provide a sound assessment of toughness,as it cannot define the relative contribution of flaws and defects from which fractureinvariably ensues. For this reason, strength and toughness can also be inversely related inceramics. By way of example, refining the grain size can limit the size of pre-existingmicrocracks, which is beneficial for strength, yet for fracture-mechanics-based toughnessmeasurements, where the test samples already contain a worst-case crack, the smaller grainsize provides less resistance to crack extension, generally by reducing the potency of anygrain bridging, which lowers the toughness.

In recent years, the development of nanostructured ceramic materials as a path to achieveenhanced strength and toughness has attracted a great deal of interest in the ceramicscommunity. However, a controversial question in composite toughening is whether the useof nanoscale reinforcements favors the strength to the detriment of the fracture toughness.Although strength and toughness may seem to many to be similar, changes in materialmicrostructure often affect the strength and toughness in very different ways and they tendto be mutually exclusive [188]. It has been claimed that composites made with suchnanoscale reinforcing materials as nanotubes, platelets, and nanofibers would have



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exceptional properties; however, results to date have been disappointing [189–191]. A goodexample here is the excitement generated by the discovery of carbon nanotubes, which haveexceptionally high strengths better than E/10 (E is Young’s modulus) [192]. However, it isstill uncertain whether such nanotubes can be harnessed into bulk structural materials thatcan utilize their high strength (and stiffness) without compromising toughness. If thecomposite material is to be used for a small-volume structure with high strength, clearly thereinforcements must also be small; moreover, as there is a lessened probability of findingdefects, small-volume reinforcements tend to be much stronger, as has been known since theearly days of research on whiskers [193]. However, from the perspective of toughening, wecould strongly argue that prime (extrinsic) toughening mechanisms such as crack deflectionand particularly crack bridging are promoted by increasing, not decreasing, reinforcementdimensions [187].

Recently, it has been shown that a possible path to combining high strength and toughness ina ceramic material is to take advantage of the transformation toughening mechanisms innanozirconia-alumina materials [194, 195]. These materials consist of a dispersion of asmall amount of tetragonal ZrO2 particles (typically around 200 nm in size) in an Al2O3matrix. The stress-induced phase transformation of metastable tetragonal grains towardmonoclinic symmetry ahead of a propagating crack leads to a significant increase of thework of fracture. For the same pre-existing defects, these composites can work at loads twotimes higher than the pure materials without delayed failure. Hardness and stability are ofprime interest in the dental field. Alumina-zirconia nanocomposites with relatively lowzirconia content (below the percolation limit, ~16 vol%) exhibit similar hardness values toalumina and are not susceptible to the hydrothermal instability observed in the case ofzirconia bioceramics that caused the in vivo degradation of zirconia femoral heads and hasresulted in the almost complete abandonment of ZrO2 as an orthopedic biomaterial [194–196].

The stabilization of the tetragonal zirconia phase in the nanocomposites is thought to arisefrom a combination of surface energy effects, constraints of the rigid matrix, and stabilizingoxide additions, e.g., yttria, such that transformation can occur locally once the constraintsare removed, in this case when the crack propagates [197–200]. It is then necessary todevelop processing methods that will allow the fabrication of materials in which the zirconiagrains have controlled submicronic size and are homogeneously dispersed in the ceramicmatrix. Colloidal procedures in which a liquid precursor of the zirconia phase is mixed withthe powders of the matrix ceramic phase present great potential. The use of a liquidprecursor allows a more intimate mixing and a more homogeneous distribution ofnanoparticles, avoiding the formation of aggregates after the heat treatment [201].

Alumina/zirconia nanocomposites offer an example of how nanotechnology offers anattractive path to the development of new implant materials but ceramics, evennanocomposite ceramics, will not replicate the unique combinations of mechanicalproperties of tooth tissues as they are, for example, much stiffer and wear-resistant. Apossibility is to develop new hybrid organic/inorganic materials whose properties willclosely match those of the tissue for which they substitute. However, the use of syntheticcomposite materials as permanent replacements for bone, which generated much excitement30–40 years ago [202], remains largely unmet due to significant challenges related tofabrication, performance, and cost. Current hybrid organic/inorganic composites havesignificant problems related mostly to their mechanical performance and their degradation invivo [203, 204]. As a result, the use of synthetic composite materials as permanentreplacements for bone is nearly nonexistent [202]. To achieve dramatic improvements in invivo performance of composites for dental implants and skeletal tissue repair, new ways ofapproaching composite design and fabrication are needed. Ideally, these materials would be



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capable of self-healing, as is the case in many biological materials. In addition, teeth have acomplex structure in which several tissues (enamel, dentin, cementum, and pulp) with verydifferent properties and structure are arranged. An ideal artificial tooth requires thecombination of several synthetic materials with prescribed properties. Will nanotechnologyalone be enough to achieve this goal?

The tissues in teeth have very specific and sophisticated hierarchical architectures, withstrength and toughening mechanisms built in at almost every dimension. Nanodimension isonly one of them but we need to develop new technologies for the fabrication of materialswith architectures designed at multiple length scales that, from the mechanical point of view,will exhibit extrinsic toughening mechanisms similar to those observed in naturalcomposites. Recent advances in the processing of bioinspired materials may open newpossibilities. In our laboratory we have pioneered a new technique, freeze casting, that canbe used to fabricate ceramic-based composites with complex architectures to assemblenovel, bioinspired hierarchical structures modeled after natural composites such as bone ornacre, striking a balance between mechanical and functional responses [205–209] (Figure 6).This technique uses the controlled freezing of ceramic suspensions to generate porousceramic scaffolds whose architecture is templated by the ice crystals. By controlling thecomposition of the suspension and the freezing conditions, it is possible to generatematerials with complex hierarchical architectures that mimic those of the inorganiccomponent of nacre at multiple length scales, from the nano to the macro levels. Thesescaffolds can subsequently be infiltrated with a second “soft” phase (e.g., polymer) toprepare composites that exhibit unique combinations of strength and toughness. Inparticular, it is possible to reach fracture resistance up to 300 times larger (in terms ofenergy) than that of their main ceramic constituents. As with natural materials, these valuesare much larger than what could be expected using the simple mixture of their components.

One of the problems for the dental application of modern technical ceramics is the difficultyassociated with the fabrication of ceramic pieces with complex shapes while maintainingaccurate dimensional control of features from the macro to the micro levels and goodmaterial properties. For example, when fabricating porous alumina dental implants to favorbone ingrowth, it has been observed that the residual unwanted surface microporosityadjacent to the gingival cuff results in an inflammatory reaction that prevents formation ofan effective biological seal and results in clinical failure [210]. The advent of the so-calledsolid-free-form fabrication technologies or rapid prototyping technologies has opened theopportunity to produce custom-designed ceramic parts directly from a computer model withdetermined shapes and porosities (Figure 7) [211–213]. These techniques produce 3Dobjects guided by a CAD file, as well as digital data produced by an imaging source such ascomputed tomography (CT) or magnetic resonance imaging (MRI), opening the door for thefabrication of custom-designed implants or crowns. Other novel processing techniques suchas injection molding or gel-casting could produce similar degrees of control.

4. Assessing implant osseointegrationA critical feature of implants intended for dental applications is their ability to promote boneformation along their surfaces. By definition, these implants need to be embedded in bone inorder to function. That function includes support for either neo-teeth or prostheses. Thebone-implant interface then becomes critical in various ways—it must secure the implantwhile distributing loads on the implant in such a way as to minimize osteolysis and implantloosening. Additionally, the interface should secure the implant quickly.

In vitro tests in simulated body fluid (SBF) are often used to assess the bioactivity ofcoatings and materials [225]. Simulated body fluid is a solution with an ionic content similar



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to human plasma. Materials able to precipitate apatite crystals on their surface duringimmersion in SBF are often considered bioactive and expected to bond to bone in vivo.However, later works have criticized the significance of the test and the procedures used asthey seem to give false positive and negative results [226]. In any case, in vitro tests in SBFseem to provide useful information on the chemical reactions that can be expected at thebiomaterials surface in vivo and can be used in combination with mechanical testing tounderstand the mechanisms of failure. Tests in cell cultures are also routinely used to assesscytotoxicity and biocompatibility as well as the effect of the materials of cell differentiation,which will be determined by examining the expression of important osteoblastic markers.However, the total lack of guidelines to link in vitro results with in vivo performance makesthe use of animal models a requirement to truly assess implant performance [227].

Two in vivo models are commonly available for assessing osseointegration viaosteoconduction, and each offers both advantages and disadvantages. The easiest model touse is the calvarial onlay [228–230]. It minimally traumatizes the animal; can be set upquickly; can be used in a wide size-range of animals, from mice to dogs; and offersstandardized data on osteoconductive bone growth into the implant. The model typicallyinvolves operatively exposing the skull, removing or elevating the periosteum, and layingthe implant directly on the bone surface. The implant can be secured to the bone, but this isusually unnecessary, since the pressure of the overlying skin helps prevent implantmigration. The implant-bone interface can be analyzed fairly easily using histology. Theprocedure and the implant do not weaken the skull. Mechanical testing of the implant can bechallenging, however, since it focuses on shear of inconveniently shaped structures. Themain disadvantage of this model relates to cost the limited size of the calvarium limits thenumber of implants that can be simultaneously tested.

The second model involves embedding implants into bone, in plug-shaped implantsembedded into same-sized holes created in either the mandible or the long bones [231]. Eachplug is gently press-fitted into the newly created hole. As new bone grows at the interfacewith the implant, it binds the implant. The implant-bone interface can then be examined viahistology and mechanical testing. Multiple implants can be placed in the same animal,although too many implants will weaken the bone and permit fractures. This model moreclosely reproduces the typical clinical situation of placing implants directly into bone.However, the procedure is more invasive than the onlay, and requires greater technical skillto place the implants properly.

5. Future trends for dental implantsDespite conflicting reports regarding the effect of ceramic coatings and micro- and/or nano-topography on the osseointegration of dental implants, the prevailing philosophy is that theymay significantly influence the bone growth and attachment to implant surfaces andultimately improve the success of dental implants and the rapid return to function (i.e.mastication). There is an urgent need for more fundamental research in this area that wouldcombine both in vitro and in vivo studies and ultimately lead to appropriate clinicalapplication. In our desire to create the perfect implant designed with nanomaterials andbioinspired approaches, we must close key gaps in our basic knowledge and create a seriesof prototype dental implants with increasing functionality. We should start by recognizingthe specific gaps in our present knowledge and then look broadly for advances that willenable us to bridge them. Specifically, we should seek to uncover the relationships linkingcomposition and materials architecture at multiple length scales with macroscopicmechanical behavior and the capability for osteogenesis. Eventually, these relationshipsmust be tested and evaluated systematically in vivo and finally in clinical studies thatmaintain the same stringent analyses and outcome measures. Though fundamental



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interactions between biomaterials and surrounding tissues may be determined, a systematicapproach must be established to confirm the relevance and consequence of these biologicresponses. Once basic strategies are defined, the knowledge could be used to design newimplant systems capable of manipulating chemistry and cellular responses down to themolecular level. With this powerful capability, implant design can be made patient-specificand customized to the biology of the patient. To reach this goal, we need to answer severalscientific questions:

• What is the optimum coating composition and thickness? Does it depend onspecific location and can it be manipulated based on the patient’s needs?

• What are the biodegradation rates of different coating materials? Are theyappropriate for their respective applications based on load-bearing?

• Is it possible to design coatings with hierarchical architectures combining goodmechanical behavior with optimum biological response?

• How do surface topography and chemistry of dental implants control cell response,in particular the differentiation of stem cells toward osteoblastic lineages? Is therean ideal number or type of the topographical features of dental implants that wouldprovide the optimum osseointegration?

• What are the best routes for the design, fabrication, and assembly of new syntheticmaterials that will mimic the properties of tooth tissue?

• Can we develop a reliable protocol to test the osseointegration and in vitroperformance of artificial teeth? How can we correlate the results from the in vitroand in vivo analyses of materials?

The answers to these questions will provide the information needed to design and engineer alibrary of dental implants for treating a diverse population of patients. This will require theintegration of interdisciplinary principles drawn from materials science, biology,computational and quantitative science, and tissue engineering, coupled with surgical insightand implant-product data gathering through comprehensive data registries. This review hassummarized some of the aspects in which nanotechnology can intervene in the design andfabrication of new scaffolds. Nanotechnology opens a new spectrum of possibilities byproviding new tools for the manipulation and characterization of matter at very smalldimensions. The goal is to build “active” implants and structures that will interact with theirsurroundings, respond to environmental changes, deliver appropriate molecules or drugs,and actively direct cellular events. The vision for the future is to move beyond thereplacement of diseased tissues to in vivo repair and regeneration. The strategies thusdeveloped should be amenable to translation into clinical therapies and treatments. To thisend, our goal is to produce materials and therapies that will bring state-of-the-art technologyto the bedside and improve the quality of life and adapt to a patient’s specific needs.

AcknowledgmentsThis work was supported by the National Institutesof Health (NIH/NIDCR) under grant No. 5R01 DE015633

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Figure 1.Successive events following the implantation of a Ti-based material. Upon exposure tooxygen in air or water, Ti and its alloys spontaneously form a stable and strongly adherentoxide film (ns timescale). After adsorption of water molecules on Ti oxide surface, the nextsequence of events is not very well understood. Immediately after implantation, implants arecoated with a fibrous connective tissue, which degrades and is replaced ideally by bone. Thethickness of fibrous connective tissue depends on the success of osseointegration. Once thecells reach the surface, their interaction takes place through the protein coatings. Adaptedfrom [13, 14, 18].



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Figure 2.Possible designs for graded coatings. The layer in contact with the substrate could beformulated to provide good adhesion and long-term stability. The top layer can vary inthickness and porosity, and can be a mixture of an inorganic and an organic component, and,if needed, be encapsulated in a bioresorbable material as either a glass (fabricated by PLD)or a polymer (by robocasting or dip-coating).



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Figure 3.Pulsed laser deposition can be used to prepare thin coatings of highly bioactive glasses thatcannot be fabricated by other methods. (a) Schematic of the PLD process. A laser beam isused to evaporate the glass into the substrate and form a film. Afterward, thermal treatmentis needed to promote adhesion. The coatings are bioactive and retain the complex glassstoichiometry. The substrate can be moved and rotated during deposition to coat 3D pieceswith complex shapes. (b) Surface of a thin glass coating deposited on Ti-6Al-4V by PLD.Thin coatings are more resistant to cracking and delamination. Because the coating hassubmicron thickness, the substrate peaks are visible in the EDS analysis.(c) Quantitativeanalysis indicates that the stoichiometry of the layer is in ±5 wt% of the original glass target.



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Figure 4.(a) Environmental scanning electron microscope image showing a cross section of a gradedglass coating after a six-month in vivo test in a pig model: the coating survived, maintainingexcellent adhesion to the implant. A corroded layer can be observed on the coating surface,providing adhesion to the bone, but the high-silica layer in contact with the alloy remainsintact, maintaining good adhesion to the metal. Preliminary tests indicate that after sixmonths, the bone matrix in the vicinity of the bone-implant interface of coated implants hadchanged to lamellar-type bone while the uncoated implants remained with woven-type bone.It is reasonable to assume that the coated device had less micromovement so that thebiomechanical integration of the device to the surrounding bone tissues was achieved earlierduring the experimental period, facilitating the transformation of woven bone matrix tolamellar type over a two- to six-month time period. (b) During indentation tests, the cracksdo not propagate at the glass/metal interface, indicating good coating adhesion achievedthrough the formation of a thin (100–150 nm) Ti5Si3 layer that can be observed using high-resolution TEM.



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Figure 5.By coating dental implants with thin films of CP glass-ceramics, and then using controlledcrystallization, nano and micro features can be obtained on the scaffold surfaces afteretching.



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Figure 6.Freeze casting, a technique that uses the microstructure of ice to template the architecture ofceramic scaffolds, can be used to produce porous lamellar materials that replicate thestructure of the inorganic component of nacre at multiple length scales [205, 207, 208, 214].These materials can be much stronger than others with similar porosity described in theliterature.



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Figure 7.Solid freeform fabrication techniques that are precise and reproducible, such as direct ink-jetprinting, robotic-assisted deposition or robocasting (in the figure), and hot-melt printing—which usually involve “building” structures layerbylayer followinga computer design, orimage sources such as MRI—can be used to fabricate custom-designed scaffolds withcomplex architectures [212, 215–224].



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