Bone graft engineering: Composite scaffolds

Chapter 6

Bone graft engineering: Composite scaffolds Jason L. Guoa, Trenton C. Piepergerdesa, Antonios G. Mikos Department of Bioengineering, Rice University, Houston, TX, United States

Chapter outline 6.1 Requirements of bone tissue engineering 6.1.1 Use of scaffolds, cells, and bioactive factors 6.1.2 Advantages of composite scaffolds for alveolar tissue engineering 6.2 Composite materials and scaffolds used for bone graft engineering 6.2.1 Synthetic polymers 6.2.2 Biological materials

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6.2.3 Ceramics 6.2.4 Scaffold fabrication methods 6.3 Composite scaffold design for the reconstruction of alveolar bone 6.3.1 Monophasic scaffolds 6.3.2 Biphasic/multiphasic scaffold design 6.3.3 Gradient scaffold design 6.4 Conclusions and additional considerations Acknowledgments References

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Abbreviations 3DP BMP BMSC DPSC ECM HAp hDPC PCL PDL PDSC

three-dimensional printing bone morphogenetic protein bone marrow-derived mesenchymal stem cell dental pulp stem cell extracellular matrix hydroxyapatite human dental pulp cell poly(ɛ-caprolactone) periodontal ligament periosteum-derived stem cell

a. These authors contributed equally to this work. Dental Implants and Bone Grafts Materials and Biological Issues. https://doi.org/10.1016/B978-0-08-102478-2.00007-6 Copyright © 2020 Elsevier Ltd. All rights reserved. 159

160  Dental implants and bone grafts materials and biological issues PEG PGA PLA PLGA TCP TGF-β

poly(ethylene glycol) poly(glycolic acid) poly(lactic acid) poly(lactic-co-glycolic acid) tricalcium phosphate transforming growth factor beta

6.1  Requirements of bone tissue engineering In minor cases, bone has an innate capacity for regeneration and can thus heal without external aid. There are cases, however, where the scope of the injury or defect is beyond the regenerative capacity of the bone, thus necessitating external intervention. Bone tissue engineering utilizes scaffolds, cells, and bioactive factors, either alone or in combination, to encourage bone regeneration in defect sites. Scaffolds can provide structural and mechanical integrity, while bioactive moieties provide cell- and tissue-specific cues to the native cells. One can also deliver exogenous cells that deposit extracellular matrix (ECM) and remodel the tissue architecture without relying upon native cells. Through various combinations of these three components, tissue engineering seeks to mimic biological, mechanical, and physical properties of the damaged tissue. This section will discuss the three main components of tissue engineering strategies as they apply to alveolar bone regeneration. Further it will focus on how these components are employed in combinations to achieve tissue regeneration in periodontal bone defects.

6.1.1  Use of scaffolds, cells, and bioactive factors Insult or injury in the maxillofacial region is often structurally complex in three dimensions (3D) and susceptible to various mechanical forces at the defect site. Any tissue engineering scaffold must therefore have mechanical stability within the defect space. This is achieved in tissue engineering through the use of materials capable of forming stable 3D structures that define and maintain a 3D space. These scaffolds can be synthetic, natural, or composite materials, with specific materials being discussed later in this chapter [1]. In general, scaffolds act to define the zone of regeneration and provide structural support to tissue as it forms within the defined zone. A requirement in scaffold design is matching the mechanical properties of the material to the tissue defect. For alveolar bone defects, this necessitates mechanically strong materials that can support the mechanical forces common to the region (e.g., mastication) [2, 3]. Another key parameter in scaffold design is the porosity of the material. Pores must be sized such that cells can infiltrate the scaffold and initiate vascularization and tissue deposition [4]. For alveolar bone, optimal pore sizes and porosities have been roughly 150–500 μm and 70%, respectively [5]. A caveat here is that the mechanical properties of the scaffold decrease as porosity and pore size increase, so one must be careful to optimize porosity for cellular infiltration and maintenance of sufficient mechanical properties [6]. Scaffolds must also display

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biocompatibility so that native cells and ECM components can interact in and around the scaffold to form new tissue. Further the scaffold must be biodegradable in order to create space for the infiltration and deposition of tissue components by native cells. Interestingly, alveolar bone formation is much quicker than long bone formation, so degradation timescales of 5–6 months have shown to be ideal [5]. Most common scaffold materials are limited in their ability to induce tissuespecific signaling, due to a lack of appropriate biological cues. In order to remedy this, tissue engineering efforts often incorporate tissue-specific, bioactive molecules such as growth factors. Here, we will consider any molecule or signal that induces phenotypic changes in receiver cells as a bioactive molecule. For alveolar tissue engineering, bone-specific growth factors are popular candidates to induce tissue formation [7]. Growth factors can bind one or more cell surface receptors to elicit many different cellular responses. Of importance to alveolar tissue engineering are transforming growth factor beta (TGF-β), bone morphogenetic proteins (BMPs), and platelet-derived growth factor, among others, as they are all key players in osseous wound healing [7]. TGF-β and its isoforms are expressed during every phase of bone wound healing and are responsible for recruiting regenerative cells, inducing regenerative signaling cascades, encouraging vascularization of healing tissue, and enhancing osteoblast function [7]. For alveolar and periodontal applications, TGF-β has improved bone healing near dental implants and enhanced integrity of regenerated periodontal ligaments, but results have been inconsistent across animal models [8, 9]. BMPs are often employed for their potent capacity for directing osteogenic differentiation of cells. Isoforms BMP-2, -4, -7, and -12 are most frequently used for alveolar bone tissue engineering and have shown significant improvements in bone volume and degree of vascularization of regenerated bone. Further, the delivery of combinations of growth factors has yielded promising results as it more closely mimics the complexities of multiple growth factor expression during alveolar bone regeneration [7, 10]. While promising, the use of purified growth factors and peptides for tissue engineering suffers from a few limitations. A finite amount of biomolecule is delivered at scaffold implantation, and these biomolecules are subject to diffusion away from the defect site as well as loss of bioactivity. This necessitates the incorporation of supraphysiological concentrations of proteins and peptides in order to achieve a therapeutic effect over timescales relevant to tissue formation (weeks to months). Further, high levels of potent growth factors, like those listed previously, can yield inflammation, tissue overgrowth, and even tumorigenic phenotypes in recipient cells, while growth factor diffusion can yield off-target effects [11]. For these reasons, controlling the spatial and temporal release profiles of these growth factors in their active form is crucial to their success. Cell delivery offers different advantages to alveolar tissue engineering interventions. First, if cell type is chosen appropriately for alveolar bone as will

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be discussed next, then delivered cells will be able to express multiple required growth factors and cellular products necessary for the multiple stages of osseous tissue healing [12]. This growth factor expression is theoretically infinite during the lifetime of the delivered cell and can exhibit dynamic temporal patterns characteristic of growth factor expression in  vivo. Cellular interventions thus achieve complexity far greater than single or even multigrowth factor delivery. Here we will highlight a few key cell types that have demonstrated efficacy in alveolar tissue engineering. Native tissue regeneration relies upon stem cells to produce local and systemic signals, recruit relevant cells to the injury/defect site, and differentiate into tissue-specific cell types. Of the many possible stem cell sources, bone marrow-derived mesenchymal stem cells (BMSCs) seem to be prime candidates for alveolar bone regeneration. BMSCs have the innate ability to differentiate into osteoblasts and chondrocytes when delivered in  vivo [13–15]. Further, these cells have demonstrated efficacy in alveolar ridge augmentation in that their delivery has been shown to increase bone volume and neovascularization in regenerated tissue when compared to the usage of acellular scaffolds alone [16, 17]. Another source of stem cells for repairing the alveolar defect therapies is the periosteum, the thin layer of regenerative components surrounding bones [18]. These periosteum-derived stem cells (PDSCs) display similar differentiation markers as BMSCs and are more easily harvested. Further, autologous transplants of PDSCs have shown successive bone regeneration and angiogenesis in animal models and clinical trials [19–22]. Other popular stem cell sources that have demonstrated efficacy for alveolar and periodontal engineering are PDL stem cells [23–26], dental pulp stem cells (DPSCs) [27–29], and adipocyte-derived stem cells [30–32]. Many groups have also combined the differentiation potential and heterotypic cell-cell signaling of multiple stem cell types to achieve more accurate mimicry of the native biological milieu [33]. Despite this, cellular therapies display many weaknesses that limit their clinical promise. First, stem cell harvesting most often yields heterogeneous populations of cells and the maintenance of differentiation capabilities of harvested stem cells remains a challenge due to factors associated with in vitro cell expansion, such as strict criteria for media composition, culture conditions, and culture duration [34]. Additionally, stem cells delivered to in vivo defects tend to have low survival and significant migration away from the defect site [35]. In periodontal injuries, multiple cell types (osteoblasts, cementoblasts, periodontal ligament cells, endothelial cells, fibroblasts, etc.) signal to one another for repair and maintenance of the periodontal unit. Thus tissue engineers must often deliver either heterogeneous cell formulations and/or heterogeneous bioactive factor formulations (e.g., decellularized ECM [36, 37], calcium phosphate doping [38, 39], and growth factors as described earlier). While they present many advantageous traits for tissue regeneration, cellular and bioactive factor therapies are ultimately dependent upon a scaffold that defines the space with sufficient biological fidelity and exhibits the appropriate

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physical and mechanical cues. In this chapter, we seek to highlight ways that tissue engineers have optimized material selection and composite scaffold design for alveolar tissue engineering.

6.1.2  Advantages of composite scaffolds for alveolar tissue engineering As mentioned, every material has benefits and drawbacks for its use in alveolar tissue engineering. When seeking to remedy injury to the alveolar bone, which resides within a multitissue unit, one is faced with a myriad of needs, including spatial differences in biological, mechanical, and physical composition. Single materials thus cannot match the complexity associated with injuries to these tissue sites. It is therefore necessary to design composite scaffolds that combine the advantages of multiple materials to better mimic the heterogeneous environment found in injuries to the periodontal region.

6.2  Composite materials and scaffolds used for bone graft engineering Scaffolds used for the tissue engineering of dental bone grafts must recapitulate the physical, mechanical, and biological properties of the native bone. Given that the scaffold in a bone graft acts as a surrogate ECM while bone regenerative processes occur post-implantation, the materials used for bone tissue engineering should mimic both the chemical and physical cues seen in the ECM of alveolar bone or surrounding tissues [40]. However, no single material or class of materials can effectively mimic all aspects of the native ECM. Thus composite materials and scaffolds have been developed for the tissue engineering of alveolar bone and its related tissues. These scaffolds often utilize combinations of synthetic and natural materials to provide a more diverse milieu of tissue regenerative cues to encapsulated and native cells in the periodontium. This section seeks to outline the unique benefits and drawbacks provided by the individual materials used in composite scaffolds for dental bone. Synthetic polymers—ranging from polyesters to polyurethanes—are one class of materials that provide highly tunable properties and good processability to fit the physical and mechanical needs of dental bone, but lack the tissue-specific biological cues for bone regeneration. Biological materials, such as ECM-derived polymers and bone growth factors, present these cues but lack the mechanical properties to support the load requirements of dental bone. Ceramics, on the other hand, have been popularly used in bone grafts and offer both chemical cues for bone regeneration and strong mechanical properties, however, they present issues of poor degradation and brittle fracture when used as singular materials [41]. Metals, while commonly used in dental implant materials, are used less frequently in tissue-engineered bone grafts due to their suboptimal degradability in  vivo, with the exception of certain metal materials such as

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Mg-based metals [42, 43]. Materials from different classes are thus combined to generate the full array of desired physicochemical and biological properties for regenerating dental bone.

6.2.1  Synthetic polymers Synthetic polymers offer highly tunable mechanical properties and degradation behavior by tuning parameters during either polymer synthesis or processing [44]. Composite materials with varying properties can thus be generated by the co-polymerization of multiple monomers, the usage of polymer blends, or the combination of synthetic polymers with other materials such as biological polymers or ceramics. Polyesters such as poly(ɛ-caprolactone) (PCL), poly(glycolic acid) (PGA), and poly(lactic acid) (PLA) are popularly used due to their degradability in vivo via hydrolysis of ester bonds and ease of co-polymerization to generate polymers with variable properties (Fig. 6.1). Polyesters such as PCL, PGA, and PLA are particularly advantageous due to their degradability by hydrolysis and ease of co-polymerization to tune resultant polymer properties. Poly(lactic-co-glycolic acid) (PLGA) is a common copolymer and polyester that has been used throughout bone tissue engineering. In dental tissue engineering, it has been used to provide mechanical support to BMP-2-loaded gelatin sponges for the treatment of alveolar bone defects in an in vivo adult beagle dog model, which produced the formation of both new alveolar bone tissue and cementum in the defect [45]. In another instance, lactic acid was copolymerized with ɛ-caprolactone to generate a similar polymer support for basic fibroblastic growth factor-loaded gelatin sponges targeting the alveolar ridge [46]. These

FIG. 6.1  Generalized chemical structures of commonly used polymers in bone tissue engineering.

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examples show a common strategy in which the synthetic polymer is used to provide structural support to an otherwise mechanically weak biological material in the tissue engineering scaffold. Other synthetic polymers such as polyurethanes have been utilized more rarely for alveolar bone tissue engineering due to their longer degradation times compared to the more highly labile bonds of polyesters. The degradability of these polymers is crucial to allow for their replacement by natural ECM over the course of tissue regeneration [44]. Wu et al. showed one case, however, in which polyurethane sponges were used as intermediate materials to generate macropores during the fabrication of a bioactive glass scaffold for the treatment of alveolar bone defects [47]. PCL, PLGA, and other polyesters can be processed by a wide range of techniques ranging from electrospinning to rapid prototyping, which allows for the creation of scaffolds with highly defined geometries, even at the millimeter scale required for reconstruction of alveolar bone [48, 49]. In one case, Li et al. 3D printed patient-specific PLGA and tricalcium phosphate (TCP) scaffolds using computed tomography scans and 3D models of the patient’s alveolar bone defect, producing anatomically shaped tissue growth [48]. Control of scaffold architecture shown here and in other examples involving synthetic polymers is particularly advantageous when designing composite grafts that target both the alveolar bone and periodontal ligament (PDL), which requires gradation of mechanical and physical properties along the transitions between tissue types [48, 49]. Vaquette et al., for instance, utilized fused deposition modeling and electrospinning to process PCL into alveolar bone and PDL compartments, respectively, with distinct mechanical properties and differential cell types encapsulated within each compartment [49]. As seen in many of the aforementioned cases, synthetic polymers are usually processed in tandem with biological materials or ceramics to generate composite scaffolds that effectively address the shortcomings of using synthetic polymers alone. For instance, both biological polymers and ceramics can provide the tissue-specific cues for alveolar bone not presented by the aforementioned synthetic polymers, which are largely bioinert [47].

6.2.2  Biological materials Naturally derived polymers, such as collagen, gelatin, and chitosan, are often used as scaffold materials or carriers for growth factor delivery to the defect site due to their inherent bioactivity and degradability in vivo. Collagen and gelatin are proteins found in the native ECM and are well suited for tissue engineering of various tissues including alveolar bone and the PDL, due to their general propensity to support the adhesion and proliferation of many cell types [44, 50]. Chitosan, on the other hand, is a polysaccharide with not only cell-adhesive properties but also antibacterial behavior that is particularly advantageous for the alveolar bone, which is by nature highly susceptible to bacterial infection [51]. An alternative to fabricating scaffolds from individual biological polymers

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is to extract and decellularize the native ECM from the alveolar bone or PDL itself, producing a mixture of ECM components that closely resembles that of the native tissue [52]. Farag et al. demonstrated one interesting approach in which they cultured PDL cells in a monolayer, decellularized the resulting cell sheets, and then characterized high recellularization potential in their PCL-supported, decellularized sheets [36]. Their findings indicate the ability of decellularized ECM to produce accurate biological cues for a given tissue [36, 52]. However, a major drawback of decellularized ECM and other natural polymers is their inability to support the mechanical loads experienced by alveolar bone (e.g., loading during mastication) necessitating the co-fabrication of ceramic or synthetic polymer components in the scaffold [44]. In one case, Peter et al. doped hybrid chitosan-gelatin scaffolds with ceramic nanoparticles to both increase the mechanical properties and induction of mineralization by osteoblast-like MG-63 cells [51]. Many other examples exist where gelatin or collagen are used in combination with polyesters such as PCL or PLGA for the reconstruction of alveolar bone and other periodontal tissues [45, 46, 53]. Aside from biological polymers, osteogenic growth factors are often critical components of composite scaffolds for the repair of alveolar bone, as they provide direct tissue-specific cues for osteogenesis. The BMP family of growth factors, BMP-2 and BMP-7 in particular, have been investigated in many instances for the regeneration of alveolar bone, with several in vivo studies showing the efficacy of scaffold-mediated BMP delivery in animal defect models [45, 53, 54]. One interesting issue that arises in the repair of whole periodontium is the necessity for strict localization of BMP and other osteogenic growth factors to the alveolar bone to prevent mineralization and bone ingrowth to surrounding nonosseous tissues. Kuboki et al., for instance, found that a multilayered scaffold with restriction of BMP to the alveolar region was necessary to prevent ankylosis and ingrowth of alveolar bone to the dentin tissue in a baboon model [54]. Thus issues of ectopic tissue growth must be addressed when designing a composite scaffold for any tissue within the periodontium, for instance, by binding or otherwise localizing tissue-specific factors to their respective layers.

6.2.3 Ceramics Ceramic materials are typically used as an additive for composite scaffolds as they provide mechanical strength and potent chemical cues for mineral deposition, while limiting the drawbacks of poor degradation and brittle behavior associated with pure ceramic scaffolds. Calcium phosphates such as TCP and hydroxyapatite (HAp) are commonly used due to their similarity to the inorganic phase of bone and have been co-processed with synthetic polymers such as PCL and PLGA to generate mineralized composite scaffolds [46, 48, 49]. These material “cocktails” can also be processed and printed as a single mixture by rapid prototyping technologies such as fused deposition modeling and lowtemperature deposition manufacturing, allowing for the creation of scaffolds

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with highly uniform mineral content [48, 49]. However, the conditions for printing or otherwise processing calcium phosphates and other ceramics are often harsh (e.g., requiring supraphysiological temperatures), so cells and biological materials must often be incorporated by other means such as post-fabrication seeding or a secondary mode of fabrication such as electrospinning [55]. Bioactive glass is another type of ceramic material that has received recent interest due to its ability to induce mineral deposition, bind and integrate with bone via formation of an intermediate apatite mineral layer, and ultimately stimulate osteoblastic differentiation [51, 56]. Mesoporous bioactive glasses, in particular, have attracted attention for use as particulate additives given their nanoscale pores provide enhanced surface area for the aforementioned mineralizing effects to occur [47, 57]. A few examples have emerged in alveolar bone tissue engineering with the usage of bioactive glasses as highly osteogenic additives. In one case, Peter et al. doped bioactive glass particles inside a composite chitosan-gelatin scaffold to improve mineral deposition, cell attachment, and cell spreading by osteoblasts [51]. In another example, Wu et al. produced strontium-doped mesoporous bioactive glass scaffolds, and found that increasing concentration of strontium, which is found as a trace element of native bone hydroxyapatite, augmented osteogenesis by encapsulated periodontal ligament cells [47]. Compared to traditional ceramic materials, these mesoporous and ion-doped bioactive glasses thus present more osteogenic cues for the repair of alveolar bone.

6.2.4  Scaffold fabrication methods A nontrivial variable for the production of scaffolds, either single material or composite, is the method of fabrication of the final construct. Different fabrication methods are compatible with different materials and provide myriad strengths and weaknesses, so the fabrication method must be selected with the end goal in mind. Here, we will describe common methods for the manufacturing of biomaterial scaffolds with an emphasis on their use in alveolar tissue engineering. Crosslinking of water-soluble polymers is a common prerequisite for fabricating stable tissue engineering constructs from a collection of individual synthetic or natural polymers. Further, polymer crosslinking can be used to create insoluble polymer networks, allowing for the creation of aqueous scaffolds such as hydrogels, which can be used under in vitro or in vivo conditions. In the case of many natural polymers, physical interactions (charge affinity, hydrophobic interactions, hydrogen bonds, etc.) are enough to provide structural stability, but do not confer sufficient mechanical strength. The use of chemical crosslinkers introduces the variable of crosslinking density that can be tuned to achieve specific scaffold properties in mechanical strength, swelling, degradation kinetics, and porosity of the resulting scaffold. In this way, scaffolds can be optimized to best serve the regeneration of the respective tissue. For example, crosslinking

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density and the corresponding degradation kinetics can be optimized to provide mechanical stability early in tissue regeneration, then allow newly formed ECM to replace the scaffold material as it degrades [58]. A critical concern for any scaffold fabrication method is the creation of tissuespecific 3D architectures. As mentioned previously, an interconnected porosity is necessary for the infiltration of cells and vasculature once the scaffold is implanted into the defect site. To achieve this, porogens can be used to create a macroporous structure. A porogen is any material that can dissolve or degrade while the overall scaffold retains its structure, leaving vacant spaces in its original locations. Popular biocompatible porogens for bone tissue engineering scaffolds include salts, glucose, and gelatin [59, 60]. Demarco et al. compared gelatin microspheres and salt as porogens for guided differentiation of dental pulp stem cells into odontoblasts [61]. They found that PLLA scaffolds made porous with salt lead to earlier odontoblastic differentiation, but PLLA scaffolds made porous by glucose microspheres increased expression of differentiation markers at later time points [61]. While simple to execute, this approach results in significant pore size and shape variation, limited interconnectivity, and limited scaffold size, as leaching the inner porogens gets more difficult as the scaffold thickness increases [62]. A scaffold fabrication method that addresses some of the aforementioned limitations is electrospinning, which begins with the solubilized polymer inside a syringe. Some distance away, a collector plate is set perpendicular to the needle and a voltage is applied from the needle to the plate. As the solution is extruded from the syringe, the charge of the polymer causes it to travel in a Taylor cone to the plate. As the polymer travels between the needle and surface, the solvent evaporates and the polymer is deposited on the plate as a continuous fiber. In this way, meshes with randomly aligned polymer fibers are formed on the charged surface. Through altering the solution extrusion rate, the needle size, and applied charge, one can precisely tune the resulting fiber diameter, porosity, and mesh thickness [63]. This method has allowed Inanç et al. to create nanofibrous PLGA meshes that encouraged the adhesion of PDL cells along with marked improvements in ECM deposition [64]. Shin et al. similarly demonstrated effective bone formation in alveolar defects with an electrospun PCL mesh seeded with BMSCs [65]. Through a simple CaP coating, Dan et al. employed an electrospun PCL mesh in a rat periodontal defect and demonstrated the formation of new alveolar bone, cementum, and PDL tissues at 4 weeks [66]. While useful for the creation of cell supportive scaffolds, these traditional scaffold fabrication methods are limited in their ability to achieve complex, homogeneous, and spatially defined architectures [67]. In recent years, threedimensional printing (3DP) techniques have proven themselves as useful tools for the high-precision manufacturing of both mono-material and composite scaffolds. 3DP allows the user to digitally design a variety of 3D architectures, allowing for customizable geometries that can match those of the defect site. Further, 3DP methods are compatible with many of the different scaffold ma-

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terials described throughout this chapter, both alone or as composite “inks,” and can be adapted to incorporate bioactive factors and cells. In simpler cases, 3DP can allow for the fabrication of highly precise and patient-specific molds in which to create scaffolds via traditional crosslinking-based methods [68]. In some cases, micro-computed tomography scans of a defect can be translated to CAD models that allow 3DP of constructs specifically tailored to the defect geometry. Most early work with single material printing employed fused deposition modeling, a technique that heats thermoplastic materials to their melting temperature then deposits them as viscous materials onto a platform in a controlled manner. That material then solidifies as it cools and the deposition head can begin adding layers. PCL lends itself well to this technique and has been extensively investigated for bone and periodontal applications. Hutmacher et al. explored how pore structure impacted the mechanical properties of these 3DP PCL scaffolds [69]. With a constant porosity set at ~60% and a fiber thickness of 1.70 mm, two different fiber alignments between layers were tested for compressive stiffness and offset yield strength. The three-angle scaffold proved to be much stiffer than the five-angle scaffold, suggesting that the 3D porous structure plays a significant role in the mechanical properties of the scaffold. The osteogenic capacity of these scaffolds was further improved by the addition of β-TCP in the print solution [70]. A similar strategy was used to print an α-TCP/HA scaffold with fiber diameters around 400 μm and a porosity nearing 60% [71]. This material demonstrated improved osteogenic signaling and degradation times, but no mechanical analysis was performed and all in vivo work was performed in nonload bearing locations. Using an indirect rapid prototyping technique where a mold was printed to caste the personalized scaffold, Schumacher et  al. explored how factors such as calcium phosphate phase composition, scaffold macroporosity, and scaffold pore geometry impacted the resulting mechanical properties [72]. When comparing material composition, the group explored pure HA, pure TCP, and a gradient of four composite ratios in between. Experiments demonstrated that a HA:TCP ratio of 60:40 achieved the greatest compressive strength that mimicked that of native bone. To achieve these strengths, however, the group employed much lower porosities and pore sizes (~20% and ~340 μm, respectively). These features would most likely prevent cell infiltration and vascularization, but no cellular assays were performed. Another group implemented similar materials but employed a laser sintering method where extreme temperature sufficient for ceramic polymerization is achieved through laser excitation [73]. Mangano et  al. were able to create a 3DP composite HA/TCP scaffold via sintered dispense-plotted assembly [73]. With this method, fiber diameters of roughly 300 μm formed the structure of a 60% porous scaffold with pore sizes averaging 370 μm. In vivo experiments demonstrated proper scaffold degradation in relation to new bone formation, direct osteogenesis, vascularization, and negligible immune response.

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Other groups have explored optimizing the 3DP material for not only the osteogenic layer, but the PDL and cementum layers as well. Lee et al. employed a well-established PCL/HA blend but varied the fiber diameter and porosity to achieve differential tissue formation [74]. Larger fibers and smaller pores showed successful cementum/dentin mimicry, while larger pores and thinner fibers more closely mimicked the fibrous and mechanically soft structure of the PDL. Guidance of specific tissue formation was confirmed by RT-PCR analysis of mRNA expression specific to each tissue type. Park et  al. attempted a 3DP approach with a different material to match PDL properties [75]. A CADguided mold was created and PGA was cast. This polymer more closely mimics the soft tissue of the PDL as compared to PCL. Through casting PGA scaffolds and combining with a PCL osteogenic scaffold, the group was able to design a composite scaffold. This modular technique allowed them to control where osteogenic tissue was deposited on the scaffold as compared to polarized fibrous tissue characteristic of the PDL. The aforementioned methods require melting and sintering temperatures far above cytotoxic levels, thus preventing the incorporation of cells or bioactive factors. Many groups have therefore explored more cytocompatible printing conditions. Xue et al. implemented a sodium alginate-gelatin hydrosol solution in which they suspended human dental pulp cells (hDPCs) and 3D structures were created using a biocompatible extrusion-printing platform [76]. The resulting scaffold was structurally sound and hDPCs exhibited an 87% survival rate through the printing process. Another group doped a printable alginate hydrogel with decellularized dentin and explored how printing parameters impact accuracy, cell viability, and odontogenic potential. Increasing the ratio of decellularized dentin to alginate yielded less viscous materials without impacting the shear thinning properties of the material. It was also elucidated that the feed rate of the printer must relate indirectly to the viscosity of the material to achieve appropriate printing resolution. Another interesting finding is that the less viscous the material, the less capable the material is of printing sharp turns or highly regular lines. Finally, the group demonstrated that the printing was compatible with cells and could induce osteogenic differentiation in vitro [77]. Overall, scaffold fabrication depends upon the selected material and the properties of the tissue of interest. As described, multiple approaches have shown efficacy towards creating scaffolds for alveolar bone and periodontal tissue engineering. Future approaches will improve the resolution of emerging methods in 3DP and the ability to multiplex scaffold materials in order to match the heterogeneous properties of dental tissue.

6.3  Composite scaffold design for the reconstruction of alveolar bone While early tissue engineering strategies have focused on the fabrication of homogeneous scaffolds for single tissue types, the regeneration of alveolar bone

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and many other tissues in the body often relies on the successful repair of multiple, co-dependent tissue types [78, 79]. For the alveolar bone, this includes other tissue types in the periodontium, namely, the gingiva, cementum, and PDL, given that trauma and degenerative disease often insult multiple tissue types within the periodontium [53]. Thus the successful tissue engineering of any single tissue component will require targeting of the distinct properties of both hard tissues (i.e., the cementum and alveolar bone) and soft tissues (i.e., the PDL and gingiva), which present distinct mechanical and biological properties [80]. The PDL in particular represents an interesting challenge in which the PDL-bone and PDL-cementum interfaces present gradated properties which have been difficult to recapitulate using traditional scaffold fabrication technologies [49, 54]. The following discussion serves to highlight how composite scaffold materials have enabled more effective biomimicry of these periodontal tissues, and furthermore how multiphasic and gradient-based scaffold design can be used for the targeting of multiple tissue types within a single scaffold.

6.3.1  Monophasic scaffolds As discussed, single materials do not capture the mechanical, physical, and biological complexity of dental or maxillofacial defects. Further, no single biomaterial candidate affords all necessary properties to induce sufficient tissue generation. For this reason, the field of dental tissue engineering has turned towards composite scaffolds to achieve combinatorial properties for more successful interventions. We will first discuss ways in which composite yet monophasic scaffolds have been investigated. Here we define monophasic as the final scaffold being uniform or homogeneous throughout while still being comprised of more than one material. Early dental interventions employed metals to replace alveolar bone because of their mechanical strength and often bioinert properties. However, it was soon elucidated that metals did not integrate well with the surrounding tissues, especially when the alveolar ridge was insufficient, nor did they achieve functional rehabilitation at desired rates [81]. To achieve greater osteogenic capacity for metal implants, calcium phosphates were employed. Bioactive ceramic glasses were developed that could covalently bind to metal surfaces and allow integration of metal implants with surrounding tissues [82, 83]. A similar approach has been employed where the osteogenic capacity of polymeric scaffolds is enhanced with the addition of calcium phosphates. Kim et al. implemented a 3DP approach where a hybrid PCL/hydroxyapatite scaffold was fabricated via a layer-by-layer apposition approach to achieve a geometry identical to the induced incisor defect [84]. This scaffold was then used to deliver growth factors stromal-derived factor-1 and BMP-7 to assist in the induction of osteogenesis. This acellular technique was able to generate an anatomically accurate tooth and periodontium via the composite approach. Davies et al. employed a similar approach where a PLGA scaffold was made more osteogenic

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via the addition of tetracalcium phosphate and dicalcium phosphate anhydrous as small particles in the pre-polymer solution [85]. The incorporation of calcium phosphates in the scaffold decreased the total scaffold area enveloped in foreign body giant cells and promoted formation of trabecular bone. Many other groups have similarly doped synthetic or natural polymeric systems with calcium phosphates to improve osteconduction with notable success [86–88]. While these monophasic composite interventions show promise in regenerating alveolar bone alone, they fall short in their ability to stimulate integration with surrounding tissue or produce growth of multiple tissue types. These monophasic approaches thus show poor feasibility for usage in a clinical scenario. Therefore the presentation of multiple material cues, and the spatial organization of these cues in a biomimetic manner, are critical when dealing with heterogeneous defect spaces such as the periodontal space.

6.3.2  Biphasic/multiphasic scaffold design Multiphasic scaffold design utilizes co-material mixtures, additives, or separately fabricated scaffold regions to produce “phases” with distinct mechanical and/or biological properties, mimicking different tissue microenvironments. Biphasic scaffolds are the simplest and most common approach in this regard, where two stratified scaffold regions are produced for the targeting of two connected tissue types. Park et al. demonstrated such an approach by printing PDL cell-encapsulated PGA columns on top of and oriented perpendicularly to a BMP-7-modified and fibroblast-encapsulated PCL region, targeting the PDL and alveolar bone, respectively (Fig. 6.2) [75]. As shown in this example, different cell types and their associated growth factors can be cultured in separate layers to support differential tissue development; in this case, PDL cells for the PDL and BMP-7 treated gingival fibroblasts for the alveolar bone. Scaffold architecture can also be tuned to produce physi-

FIG. 6.2  Bilayered, 3D-printed scaffold utilizing macroporous PCL for the alveolar bone layer and columnar PGA for the PDL layer. (Reproduced with permission from Park CH, Rios HF, Jin Q, Bland ME, Flanagan CL, Hollister SJ, et al. Biomimetic hybrid scaffolds for engineering human tooth-ligament interfaces. Biomaterials 2010;31:5945–52.)

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cal cues in the scaffold for cell differentiation along different tissue lineages. The emergence of rapid prototyping technologies has allowed for greater control of these physical scaffold parameters, which is particularly useful for the fabrication of multilayered scaffolds targeting heterogeneous tissues [74, 84, 89]. In one notable example of multilayered design, Lee et al. produced a triphasic scaffold for alveolar bone, the PDL, and the cementum-dentin interface via the layer-by-layer printing of three scaffold sections with distinct microchannel diameter [74]. The seeding of alveolar bone stem cells, PDL stem cells, and DPSCs within the microchannels of each respective layer, along with growth factor formulations, then produced formation of mineralized bone tissue, perpendicular PDL fibers, and dentin/cementum-like tissue in vitro. Furthermore their triphasic scaffold design was able to generate phenotypically consistent tissue formation in vivo with DPSCs as the sole encapsulated cell type, indicating the strong potential for scaffold physical cues and bioactive factor presentation to direct cell differentiation along appropriate lineages even in tissues as heterogeneous as the periodontium [74]. The usage of ceramic additives can also provide chemical cues for the separation of hard and soft tissue layers in the periodontium. One example of this approach is to incorporate ceramic phosphates such as TCP in an alveolar bone or cementum-targeted layer while leaving a PDL or gingiva-targeted layer absent of these additives [49, 78]. The power of growth factors alone to generate multiphasic tissue development in the periodontium cannot be underestimated either. This was exemplified by Kim et al. in a study where an acellular, rapid prototyped scaffold with distinct areas of BMP-7 and stromal-derived factor-1 coating was able to produce stratified alveolar bone and PDL development, respectively, in an in vivo rat incisor model [84]. Rasperini et al. pursued another acellular approach in which a 3DP, patient-specific PCL scaffold was coated in platelet-derived growth factor BB and implanted in a human patient for repair of the periodontium, albeit with more modest tissue regeneration outcomes (Fig. 6.3) [90]. Acellular strategies such as these show promise due to their ability to avoid the issues of poor donor site availability and potential immunogenicity associated with the usage of cell-encapsulated scaffolds, issues especially pertinent to the small-scale tissues found in the periodontium [78, 84]. Scaffold physical cues, ceramic additives, and growth factor delivery may thus present a cellfree toolbox for producing stratified tissue development of the periodontium, bypassing certain translational and regulatory barriers to the clinical usage of periodontal tissue engineering therapies.

6.3.3  Gradient scaffold design After multiphasic scaffold development, the next step in replicating the heterogeneous microenvironment of the periodontium lies in producing material gradients that mimic the varying physical, mechanical, and biological properties

174  Dental implants and bone grafts materials and biological issues

FIG. 6.3  3D printing of acellular, patient-specific PCL scaffold to support regeneration of PDL and alveolar bone. A cross-section of the composite scaffold (bottom left) shows distinct structures for PDL and bone repair in red and blue, respectively.

of the native tissue unit. The soft-to-hard tissue transition found in the native PDL, for instance, presents gradients of not only ECM composition but also cell phenotype and physical ECM architecture [91]. Along these lines, rapid prototyping technologies such as inkjet printing and extrusion have enabled the production of spatial gradients of scaffold materials, ceramic and growth factor additives, and even cell phenotypes, typically via the printing of multiple “inks”

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in tandem [89, 92]. Numerous examples in ligament and bone tissue engineering have shown that the usage of additives such as HAp nanoparticles or osteogenic factors such as BMP-2 can successfully induce the development of mineral and cell phenotype gradients in tissue, including mimicry of the soft-to-hard tissue transition [91]. Far fewer cases exist, however, where such a gradient design approach has been applied to the periodontium, despite its nature as a collection of heterogeneous and co-dependent tissue types. In one example, Ma et al. generated a hydrogel gradient of varying gelatin methacrylate and poly(ethylene glycol) (PEG) dimethacrylate composition to screen the effects of scaffold composition on the viability and spreading of PDL stem cells [93]. While they found that increasing amounts of PEG decreased cell viability and spreading area, it is unclear how the resultant material or mechanical properties affected cell behavior or whether these differences in material composition could have affected the differentiation potential of the PDL stem cells [93]. Thus further studies are needed to elucidate how gradients of scaffold composition or ceramic/growth factor presentation can affect the development of periodontal tissues such as the PDL or alveolar bone.

6.4  Conclusions and additional considerations While many different types of materials and composite scaffolds have been designed for the repair of alveolar bone and its associated tissues, the full collection of tissues in the periodontal complex has yet to be accounted for in any single composite scaffold design. Challenges remain, in particular, with recapitulating the gradient properties of tissues such as the PDL, an issue especially relevant to the PDL’s interface with alveolar bone and cementum. Future work can thus take advantage of advances in scaffold fabrication technology such as rapid prototyping to design multiphasic and gradient scaffolds with a higher level of complexity in mechanical and architectural properties, to better mimic the heterogeneous microenvironment of periodontal tissues. On an additional note, current alveolar bone tissue engineering strategies have by large neglected to address the relatively high microbial content of this tissue. Given that microbial infection can not only produce initial degeneration of the alveolar bone but also compromise the ability of exogenous scaffolds to produce effective tissue repair, it is evident that antibacterial therapy will be an important component of future tissue engineering strategies. Current approaches have utilized polymers such as chitosan with inherent antibacterial properties as scaffold materials, but have not yet characterized the antibacterial effects these materials or incorporated any active antimicrobial agents. Future development may thus involve the encapsulation and delivery of bactericidal drugs in alveolar bone scaffolds in the same manner that growth factors are currently delivered (e.g., by loading of drugs within a natural polymer such as gelatin). Ultimately, composite tissue engineering scaffolds have shown great promise for the generation of alveolar bone and other supporting tissues in the

176  Dental implants and bone grafts materials and biological issues

periodontal complex. The translation of these tissue engineering strategies to clinical therapies will require the traversal of barriers stemming first from the heterogeneous microenvironment of these tissues and second from the necessity of minimizing risk of infection in the dental environment. Rapid prototyping technologies, in particular, have provided potential tools for addressing the former, while the development of antibacterial and/or acellular composite scaffolds may help address the latter.

Acknowledgments We acknowledge support towards the development of biomaterials for tissue engineering by the National Institutes of Health (R01 AR068073) and the Army, Navy, NIH, Air Force, VA and Health Affairs to support the AFIRM II effort, under Award No. W81XWH-14-2-0004.

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