Natural polymers for bone repair

Natural polymers for bone repair

8

GB Ramírez Rodríguez1, TMF Patrício2, JM Delgado López3 1University of Insubria, Department of Science and High Technology (DiSAT), Como, Italy; 2IDI Group from Vangest, Complexo Industrial VANGEST, Marinha Grande, Portugal; 3Departamento de Química Inorgánica, Facultad de Ciencias, Universidad de Granada, Granada, Spain

8.1  Introduction and overview Bone is a dynamic tissue, in a continuous process of remodeling, that maintains the structural integrity of the skeleton [1]. However, bone is not capable of self-repairing large traumatic injuries due to degenerative diseases, congenital deformities, tumors, or postoperative defects [2]. Therefore clinical treatments involving the use of bone grafts for augmenting or stimulating the formation of new bone in defective areas has become essential [3,4]. The number of procedures requiring bone substitutes is increasing worldwide and will continue increasing according to the expected aging of the world population [5]. The high importance of bone tissue engineering (BTE) is made evident by the 2.2 million bone-grafting operations per year worldwide, making bone the second most transplanted organ after blood [6]. Autografts, allografts, and xenografts were the first therapeutic approaches used in bone tissue replacement. Autografts (autologous bone) are considered the gold standard, supplying growth factors, cells, and a mechanical support for tissue’s structure and so providing optimal osteoinductive, osteoconductive, and osteogenic properties [7,8]. Nevertheless, its use is limited due to the elevated fracture rates, donor site morbidity, higher care cost, and shortage of supply [6]. Allografts and xenografts have been proposed as suitable alternatives to BTE but they show poor efficient incorporation and high potential for immune rejection or pathogen transmission [9–11]. To overcome these shortcomings, a plethora of synthetic bone substitutes made of metals, polymers, ceramics, or their combinations have been designed and explored to restore bone function and regenerate the damaged hard tissue [12–17]. These synthetic scaffolds can exhibit satisfactory mechanical properties, biocompatibility, and easy processability. However, in comparison to bone autografts and allografts, they offer poor integration in the host tissue, undesired ion release (metallic scaffolds), weakness (polymeric scaffolds) [18], friability and brittleness (ceramic scaffolds) [19], and uncontrollable degradability (composites) [20]. Mimicking the biochemical and biophysical cues of bone extracellular matrix (ECM) is an interesting strategy to mitigate the adverse effects of the synthetic bone substitutes favoring cell attachment, proliferation, differentiation, and neotissue formation. Nonetheless, bone has a hierarchic structure with many levels of organization (Fig. 8.1) that make its emulation a major challenge in BTE [21]. Different suprafibrillar Bone Repair Biomaterials. https://doi.org/10.1016/B978-0-08-102451-5.00008-1 Copyright © 2019 Elsevier Ltd. All rights reserved.

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1.5 nm

c axis 50 nm ¥ 25 nm ¥ 3 nm

Collagen fibril Osteons and haversian canals

Compact bone

300 nm

67 nm

Tropocollagen triple helix

2.86 nm

Collagen molecule

HA nanocrystal Osteons 100 µm

Spongy bone

Bone Repair Biomaterials

Figure 8.1  Hierarchical structure of bone from nano- to macroscale level. HA, Hydroxyapatite. Reprinted from ref Wegst UGK, Bai H, Saiz E, Tomsia AP, Ritchie RO. Bioinspired structural materials. Nat Mater 2014;14:23 with permission of Springer Nature.

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arrangements occur and coexist, forming dense structural hierarchies from the nanoscopic to the macroscopic length scales (Fig. 8.1). At the macroscopic level, human long bone is generally composed of compact (outer) and spongy (inner) bone with a complex vascular network responsible for nutrient, waste, and gas transport (Fig. 8.1). The building block of such a complex architecture at the atomic–nanometer length scales is the mineralized collagen fibril (Fig. 8.1). Collagen triple helices are secreted from the cells and self-assembled at the molecular level in a periodic staggered array of fibrils exhibiting a characteristic banding pattern of 67 nm, where a densely packed 27 nm-long region (the so-called overlap zone) alternates with the less dense 40 nm-long gap zone (Fig. 8.1). This organic matrix acts as a template where mineral deposition occurs both inside (intrafibrillar mineralization) and outside (interfibrillar mineralization) the collagen fibrils through apparently distinct events [22]. This mineral phase consisting of poor crystalline hydroxyapatite is the main component of bone comprising 69–80 wt% (hydroxyapatite, Ca10(PO4)6(OH)2), whereas the organic matrix (type I collagen) encloses 17–20 wt%. Minor bone components are the noncollagenous proteins (e.g., sialoprotein, osteonectin, osteopontin) and water [23,24]. Natural polymers were the first biodegradable biomaterials used clinically due to their similarities with bone ECM [26]. Compared to conventional synthetic materials, natural polymers exhibit superior chemical versatility, excellent biodegradability, and improved biological performance (i.e., no toxicity, no immunological reactions, and intrinsic osteogenic capacities), which favor cell attachment and promote chemotactic responses [27,28]. In this chapter, the most commonly used natural polymers for orthopedic applications are described highlighting their main advantages and weaknesses (Table 8.1). Then, special emphasis is placed on their possible formulations (i.e., scaffold, hydrogel, and nanoparticle) and physicochemical and biological stimulus applied to synthetic biomaterials to induce in vivo bone regeneration and restore bone original functionality. Current ongoing technologies to create customized 3D structures for bone applications, such as 3D printing, are presented at the end of the chapter.

8.2  Natural polymers 8.2.1   Collagen Collagen has been widely used as a biomaterial since 1881 due to its inherent biocompatibility and biodegradability [29,30]. Collagen is the most abundant protein in mammals and it is the basic building block of ECM of different native tissues (i.e., bone, tooth dentine, skin, tendon, arteries, and cartilage). Twenty-eight different types of collagen have been identified in vertebrates together with many other collagen-like proteins [31]. Among them, type I collagen is the most widely occurring fibrillar collagen, being found in skin, tendon, bone, cornea, lung, and vasculature [29]. A type I collagen molecule consists of a long central triple-helical region of two α1 and one α2 polypeptide chains. Each polypeptide chain comprises a repeating glycine (Gly)-X-Y triplet with glycyl residues occupying every third position, and proline (Pro) and 4-hydroxyproline (Hyp) occupying the X and Y positions, respectively [29]. The three polypeptide chains are intracellular, assembled into procollagen, a triple-helical structure containing large

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Table 8.1  Summary of the most representative natural polymers used as building materials for bone repair Polymer

Source

Advantages

Drawbacks

Main formulations

Collagen

Ovine, porcine, equine, and bovine

Biocompatible, biodegradable, hydrophilic, cell-binding properties, low antigenicity

Polymeric and mineralized scaffolds

Gelatin

Denaturation and hydrolysis of natural collagen

Silk

Lepidoptera larvae, some arachnids, and various flies Brown seaweed

Biocompatible, biodegradable, cell-binding properties, low cost, gelling properties Biocompatible, biodegradable, flexible processability, high mechanical strength, thermally stable Biocompatible, biodegradable, relatively low cost, easy gelation by ionic cross-linking, and easy chemical modification with adhesion RGD ligands Biodegradable, biocompatible, high mechanical performance Bioactive, biocompatible, biodegradable, antibacterial and nonimmunogenic properties, ability for cell ingrowth Biocompatible, biodegradable, low cost

Relatively weak mechanical stiffness, low antigenicity, transmission of pathogens, induces allergic reactions Low mechanical properties, fast degradation rate Slow degradation rate (2–4 years)

Polymeric and composite scaffolds

Poor cell–material interaction due to inherent lack of cell adhesivity and low protein adsorption Low cell-binding properties

Hydrogel and micro-/ nanosphere

Alginate

Cellulose

Starch

Relatively weak mechanical strength and stability Lack of processability, low surface area

Reinforcement of composite scaffolds Sponges, composite scaffolds, hydrogels, and micro-/nanosphere Composite scaffolds

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Chitosan

Wood, plants, tunicates, and algae Exoskeleton of crustaceans and mollusks, insect cuticles, and fungi Corn, potato, wheat, and tapioca

Hydrogel and micro-/ nanosphere

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N- and C-terminal domains [32]. N- and C-propeptides are cleaved by metalloproteinase enzymes generating collagen that spontaneously self-assembles into cross-striated microfibrils (with 3–5 nm in diameter, Fig. 8.2(a))  [32]. These collagen microfibrils consist of five 1D-staggered collagen molecules that laterally pack in a quasihexagonal lattice forming a rope-like 3D structure [33]. Then, collagen microfibrils self-assemble and covalently cross-link into collagen fibers. Collagen fibers show a characteristic banding pattern of 67 nm, where the densely packed overlap region alternates with the less dense gap region (Fig. 8.2(a)) [22,34,35]. This banding pattern can be directly observed by high-resolution imaging techniques such as transmission electron microscopy (Fig. 8.2b and c) and atomic force microscopy (Fig. 8.2(d)) [36]. At the macroscale, collagen

(a)

(i) Fibrillar collagen Procollagen

Collagen

Procollagen peptidase

Fibril formation Microfibril

Stained

Aldol and aldol-histidine cross-links via lysyl oxidase

Regular patterning

(b)

(c)

Fibril Fiber

Fiber bundle (subfascicle)

(d)

0.5

1.0

1.5

2.0

2.5

µm

Figure 8.2  (a) Sketch of the collagen assembly into fibers. After procollagen is secreted into the extracellular matrix, N- and C-terminals are removed by collagen type-specific metalloproteinase enzymes leading to spontaneous assembly of collagen microfibrils. Then, collagen microfibrils self-assemble and cross-link into collagen fibers. (b and c) Transmission electron microscopy and scanning transmission electron microscopy images of a self-assembled collagen fiber, respectively. (d) 3D atomic force microscopy image of collagen fiber. Adapted and reprinted from ref Mouw JK, Ou G, Weaver VM. Extracellular matrix assembly: a multiscale deconstruction. 2014;15:771 with permission of Nature Reviews Molecular Cell Biology (Springer Nature).

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fibers organize following a specific pattern depending on the tissue. The entropy-driven self-assembly of collagen is strongly affected by the nature of collagen monomers and it is very sensitive to environmental factors (i.e., temperature, ionic strength, and pH) [29,37]. Collagen fibrillogenesis, triggering to collagen fibers similar to those of native bone, has been shown to be favored at pH values close to the isoelectric point of collagen [38]. However, pH values far away from the isoelectric point induce the lack of the periodic D-band or even the inhibition of fibril formation [39,40]. The ideal conditions for in vitro collagen assembly are: pH range between 5.0 and 8.5, ionic strength from 0.1 to 0.8, and temperatures between 15 and 37°C [39,41,42]. Collagen has become the ideal biomaterial due to its abundance, biocompatibility and absorbability in the body, hydrophilicity, easy processing, and low antigenicity [23,43]. The major natural collagen sources used for biomedical applications have been ovine, porcine, equine, and bovine [44]. Marine collagen, extracted from catfish, silver carp, and marine sponge, among others, is a cheap alternative, although its use for research and clinical usage is hindered due to its low denaturation temperatures [45,46]. Acid-solubilized (telocollagen), pepsin-solubilized (atelocollagen), fibrillar, gels, powder, sheet, sponges, and matrices are some of the collagen formulations commercially available [30]. As a function of the extraction treatment, they can be grouped as: (1) decellularized collagen matrix that maintains the original tissue properties and ECM structure; or (2) isolated and purified collagen molecules that pave the way to modulate and design materials that closely resemble the target tissue [30]. Another interesting alternative, which is commercially available, is the recombinant collagen protein (RCP) obtained through protein expression in mammalian, insect, and yeast cells [47]. Two commercial brands can be found: Cellnest and Fibrogen [48–50]. In contrast to natural collagen sources, their chemical composition and molecular weight can be controlled, they can be engineered with specific amino acid sequences enabling specific cell signaling, and they can be produced at large scale for industrial exploitation. Moreover, they exhibit a low risk associated with infectious diseases transmitted from pathogens [47]. Unlike collagen derived from animal sources, RCP has a very low tendency to form fibrils, which could be related to the lack of hydroxyproline residues [51]. Various formulations based on RCPs have been used as drug delivery platforms, scaffolds, or hydrogels, providing good cell response and osteoinductive properties [52–57]. Despite this, natural collagen encompasses the international biomaterial market like no other supplier. For instance, collagen-based hemostats (e.g., Sulzer-Spines Tech, CoStasiss Surgical Hemostat, and Floseals); collagen-based skin substitutes (e.g., TransCyte, Orcel and Apligraf, Integra Dermal Regeneration Template); collagen-based wound dressings (e.g., Biobrane and Alloderm Promogran); and collagen-based scaffolds (e.g., Duragen and Collagraft) are used for tissue repair [58].

8.2.2  Gelatin Gelatin (or gelatine) is a natural protein resulting from the denaturation and hydrolysis of collagen. It has a similar amino acid sequence to its collagenous precursor,

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although the break of the hydrogen and covalent bonds during collagen denaturation causes destabilization of the collagen triple helix triggering to a random coil structure [59]. Taking advantage of the gelling capability and viscoelastic properties, gelatin has been widely used in the food (e.g., emulsifiers, foaming agents, colloid stabilizers, fining agents, biodegradable packaging materials, and microencapsulating agents), pharmaceutical (e.g., capsules, ointments, cosmetics, tablet coatings, and emulsions), photographic, and cosmetic industries [60,61]. Gelatin is manufactured through heat treatment of collagen from different biological sources such as bovine or porcine skin, bone, or tendon. If the pretreatment procedure is in acidic conditions, type A gelatin (with isoelectric point ranging from 7 to 9) is obtained, whereas under alkaline conditions, type B gelatin (with isoelectric point of 5) is manufactured [61]. Marine collagen sources (e.g., fish skins, bones and fins, sea urchin, and jellyfish) have also been proposed for gelatin production [59], paying special attention to extraction from by-products of the fish industry as an interesting renewable material [62]. Gelatin is soluble in aqueous solution at temperatures above 40°C and turns into hydrogel through a sol–gel transition by cooling the solution below 30°C. The gelation process involves the transition from random coil to triple-helix structures, stabilized through intermolecular hydrogen bonding between amine groups of glycine and carbonyl groups of proline, creating an interconnected 3D network [63]. Gel strength and thermostability are largely dependent on the amino acid sequence and the molecular weight distribution, which result mainly from the gelatin source and processing conditions [62]. For instance, gelatin from fish has lower proline and hydroxyproline content (∼20%) than mammalian gelatins (∼30%) resulting in a reduction of the gel strength and the gelling and melting temperatures (5–10°C) [64,65]. Gelatin has been largely used in tissue engineering and regenerative medicine due to its biocompatibility, biodegradability, low cost, gelling properties, and large content of arginine-glycine-aspartic acid motifs (RGD) that act as a cell-binding site [66]. For instance, Gelfoam, a gelatin-based sponge, has been used in surgical procedures as a wound dressing to control bleeding, and it has been tested as a graft material in cardiovascular application and for bone repair [67]. Nonetheless, the use of gelatin-based scaffolds for bone repair is hampered due to its low mechanical strength and fast degradation rate under physiological conditions [59]. Due to the large number of functional groups in the gelatin backbone for coupling with crosslinkers and targeting ligands, gelatin has become a potential carrier system (i.e., hydrogel, microparticles, and nanoparticles) for controlled drug delivery (Section 8.3.2.2) [66,68,69]. For instance, biodegradable gelatin hydrogels have been used as controlled-release devices for a variety of growth factors known to enhance bone formation [70].

8.2.3   Silk Silk is a natural fibrillar protein produced by arthropods (silkworms as Lepidoptera larvae; arachnids as spiders, mites, and some scorpions; and some flies). There are two varieties of silk: mulberry silk (e.g., Bombyx mori) and nonmulberry silk (i.e., referred

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to as wild silk, with varieties such as tasar, eri, and muga). Due to its smooth texture, luster, and strength, silks from natural silkworms have been widely applied in the textile industry [71–73]. One of the most investigated mulberry silkworms is the crude silk fibroin from the silk glands of mature B. mori larvae that is composed of two proteins: fibroin and sericin (Fig. 8.3(a)). Fibroin is the structural protein of silk fibers (about 70%–80%) and sericin is the water-soluble glue-like protein that covers fibroin (20%–30%). Silk fibroin consists of a hydrophobic heavy chain (H, 325–395 kDa), a hydrophilic light chain (L, 25–26 kDa), and fibrohexamerin or P25 (30 kDa) [28,71,72,74]. The heavy chain is composed of glycine (46%), alanine (30%), and serine (12%), which are organized in a highly repetitive sequence (primary structure) containing 12 dipeptides (12 GX)n, X being 65% alanine, 23% serine, and 10% tyrosine. The nonfibrous L-chains

(a)

Protein coat

Fibroin

β-sheet crystalline domain

Sericin

Degumming

Amorphous domain

(b)

(c)

N-terminus

Heavy chain (H-chain)

(d)

C-terminus

Hydrophobic repetitive domains Hydrophilic repetitive domain

Figure 8.3  (a) Raw silk fiber composed of fibroin and sericin. (b) Sericin removed by degumming and β-sheet crystallite embedded in the amorphous matrix of silk fibroin fibers. (c) Silk fibroin heavy chain (H-chain) with repetitive hydrophobic and hydrophilic domains. (d) Atomic force microscopy image of fibrils from silkworm Bombyx mori (scale bar: 150 nm). The white arrow indicates the alignment of the fibers. Adapted and reprinted from ref Du N, Yang Z, Liu XY, Li Y, Xu HY. Structural origin of the strain-Hardening of spider silk. Adv Funct Mater 2011;21:772–778; Jao D, Mou X, Hu X. Tissue regeneration: a silk road. J Funct Biomater 2016;7:22 with permission MDPI AG and InTech (INTECH).

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consist of valine, isoleucine, and leucine. The P25 subunit, which is a glycoprotein composed of the Asn-linked oligosaccharide chain, promotes the solubility of fibroin by forming quaternary complexes [28,72,74]. Moreover, silk fibers are composed of nanofibrils with a diameter of about 20 nm (Fig. 8.3(d)) [75]. Mulberry silk proteins comprise various combinations and display four different structural components: (1) elastic β-spirals, (2) crystalline β-sheets rich in alanine (denominated silk II), (3) tight amino acid repeats forming α-helices (called silk I), and (4) spacer regions [76]. The presence of repeated sequences of alanine and glycine into antiparallel β-sheets increases the mechanical strength of the silk fiber. Moreover, the β-sheets induce cross-linking within the proteins through strong hydrogen bonds and van der Waals interactions enhancing fiber strength [77]. In fact, silk shows interesting mechanical properties (i.e., 0.6 GPa strength, 0.8 extensibility, and 70 MJ/m3 toughness), whereas its degradation rate can be controlled by secondary structure, molecular weight, crystallinity, and the amount of β-sheets [72,78]. After removal of immunogenic sericin, silk shows interesting properties for biomedical applications, such as excellent biocompatibility, biodegradability, and minimal immunogenic reactions. Adequate topographic and morphological properties for diverse cellular interactions and good cell adhesion have been reported for biomaterials based on silk fibroin. Indeed, the Food and Drug Administration has approved various medical products composed of silk fibroin for clinical practice [28,72,74,77,79–81]. Silk biomaterials are easily modeled to be applied as cell-based scaffolds and as micro- and/or nanocarriers for delivering bone growth factors, therapeutic molecules, or drugs enhancing bone tissue regeneration [72]. Moreover, silk fibroin is enriched with amino and carboxylic groups that promote the interaction with bioactive molecules, or can induce and guide calcium phosphate precipitation [72,82]. One of the nonmulberry silk fibroins (i.e., Antheraea mylitta), which contains RGD motifs, has shown to improve cellular response with respect to the native mulberry silk fibroin [72].

8.2.4   Alginate Alginate is the most abundant marine natural polymer. It is an anionic polysaccharide extracted from brown seaweed or biosynthesized by genetically modified bacteria (e.g., Azotobacter and Pseudomonas species). The latter strategy allows tailoring of the chemical composition and physical properties [83]. Alginates are linear copolymers composed of β-(1–4)-linked d-mannuronic acid (M) and α-(1–4)-linked lguluronic acid (G) residues, which are covalently linked together in different sequences (Fig. 8.4). The blocks consist of sequential G residues (GGGGGG, G-blocks), consecutive M residues (MMMMMM, M-blocks), and regions of alternating M and G residues (GMGMGM) (Fig. 8.4) [83]. The content and length of M and G chains and consequently their sequential distribution and molecular weight vary depending on the extraction source. These parameters are in turn critical factors affecting the physical properties of alginate gels [83]. The most attractive property of alginate is the ionic cross-linking that takes place when divalent cations (e.g., Ca2+ or Ba2+) bind two deprotonated carboxylate groups and two hydroxyl groups of G-units of adjacent

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Bone Repair Biomaterials COO– O

OH O

–OOC

O

O

COO–

G

G

HO

O

–OOC M

OH

O

OH

G –OOC

OH

O

OH O O

HO

M

M

OH –OOC

HO O

O HO

O

O

O

COO–

OH

OH

OH

OH

G

O HO

O

O

O OH

COO–

OH

–OOC

HO O

O

O HO

O OH

O

COO– M

G

M

Figure 8.4  Chemical structure of alginate showing the G-blocks, M-blocks, and alternating M and G residues (see main text for further details). Reprinted from ref Lee KY, Mooney DJ. Alginate: properties and biomedical applications. Prog Polym Sci 2012;37:106–126 with permission of Elsevier.

alginate chains, creating ionic interchain bridges and prompting the gelation of aqueous alginate solutions. Nonetheless, one critical drawback of these ionically crosslinked alginate gels is the limited long-term stability under physiological conditions, likely due to the ionic exchange with the surrounding medium [84]. To overcome the low stability of alginate gels under physiological conditions, covalent cross-linking has been widely investigated for greater control over stability and mechanical and swelling properties of the gels (see Section 8.3.2.1). Alginate has been used as a thickening and gelling agent and colloidal stabilizer in the food and beverage industry. In biomedicine, the use of alginate involves several advantages in comparison to other biomaterials, including high biocompatibility, low toxicity, relative low cost, easy gelation by ionic cross-linking, and easy chemical modification with adhesion RGD ligands. Therefore alginate hydrogels have been largely used in tissue engineering for: (1) wound healing, maintaining a physiologically moist microenvironment and minimizing bacterial infection; (2) the controlled release of drugs and/or growth factors (e.g., BMP-2, BMP-7, vascular endothelial growth factor [VEGF]); and (3) cell transplantation to induce tissue regeneration by means of minimally invasive surgical procedures [83,85].

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8.2.5   Cellulose Cellulose is the most abundant biopolymer in nature. This polysaccharide is the main reinforcement of plants and trees giving them functionality, flexibility, and high mechanical performance because of its hierarchical structure (Fig. 8.5(a)) [86]. Cellulose has been used as an engineering material for centuries and its use continues nowadays for a large number of applications in the paper, cardboard, cosmetics, and textile industries [86–88]. Cellulose is a homopolysaccharide mainly composed of two anhydroglucose rings (C6H10O5)n linked by (1 → 4) glycosidic bonds (Fig. 8.5(b)). The degree of polymerization of cellulose chains (n) ranges from 10,000 (wood cellulose) to 15,000 (native cotton cellulose) glucopyranose units [89]. Cellulose displays six different polymorphs (I, II, IIII, IIIII, IVI, and IVII) that can be interconverted among them. Native cellulose or cellulose I is the form found in nature [89]. Cellulose chains are interconnected resulting in a parallel stacking to form microfibrils through van der Waals and intermolecular hydrogen bonds between hydroxyl groups and oxygens of adjacent molecules. These microfibrils have disordered (amorphous) regions, which are regularly distributed along the microfibrils, alternating with highly ordered (crystalline) regions (Fig. 8.5(c)). In the crystalline regions, cellulose chains are closely packed together by strong and highly intricate intra- and intermolecular hydrogen-bond networks [86]. Wood, plants, tunicates, and algae are the most relevant cellulose sources, although in the last few decades an increasing interest in cellulose produced by bacteria such as Gluconacetobacter xylinus has emerged [86]. The cellulose source as well as its extraction process (i.e., mechanical treatment, acid hydrolysis, or enzymatic hydrolysis) determines the type of cellulose and consequently its physicochemical properties. There is a wide variety of cellulose-based materials: wood fiber, plant fiber, microcrystalline cellulose, microfibrillated cellulose, nanofibrillated cellulose, cellulose nanocrystals (CNCs), tunicate cellulose nanocrystals (t-CNCs), algae cellulose particles, and bacterial cellulose particles [86]. Among them, CNCs (including nanocrystalline cellulose, cellulose whiskers, cellulose nanowhiskers, and cellulose microcrystals) have attracted considerable attention in material science due to their unique mechanical properties (theoretical Young’s modulus value along the chain axis similar to Kevlar [90] and elastic modulus over 100 GPa) [91]. CNCs also present other advantages such as low density, low energy consumption, inherent renewability, biodegradability, and biocompatibility [92]. Hence CNCs have been widely applied as a nanoreinforcement for polymer matrices to develop enhanced biocomposites with potential use as wound dressings or bone substitutes. Nanocellulose-based composites have been used for culturing osteoblasts and chondroblasts, but studies on nanocellulose for bone tissue regeneration are still at the fundamental stage [93,94]. Bacterial cellulose has also been applied to tissue engineering as artificial skin for humans with extensive burns, artificial blood vessels for microsurgery, scaffolds for tissue engineering of cartilage, and wound dressings [95].

8.2.6   Chitosan Chitosan is the second most common natural polysaccharide after cellulose. It can be found in the exoskeleton of crustaceans (e.g., shrimps, crabs, and lobsters) and

210

(a)

Tree

Transverse section

Growth ring

Cellular structure

(b) HO C4

m

HO mm

cm

C6 C5 C3

Fibril structure Microfibril

Fibril-matrix structure

Cell wall structure S3 S2 S1 P ML

Amorphous

Elementary Fibrils

25 µm 10 nm

C5

O OH

C1

C4

HO

C6

C2

C1 O

O5 O

n

(c) Cellulose chains

Crystalline

1 nm

C2

OH C3

1→4

500 µm

Cellulose

HO3

O5

100 nm

Disordered region

Crystalline regions

300 nm

Bone Repair Biomaterials

Figure 8.5  (a) Hierarchical structure of the tree, starting from the cellulose molecule; (b) chemical structure of cellulose, showing the directionality of the 1 → 4 linkage and intrachain hydrogen bonding (dotted line); and (c) idealized cellulose microfibril showing one of the suggested configurations of the crystalline and amorphous regions. Reprinted from ref Moon RJ, Martini A, Nairn J, Simonsen J, Youngblood J. Cellulose nanomaterials review: structure, properties and nanocomposites. Chem Soc Rev 2011; 40:3941–3994 with permission of The Royal Society of Chemistry.

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mollusks, insect cuticles, and fungi. The use of chitosan covers multiple industrial applications in pharmacy, textiles, paper, food, agriculture, biotechnology, and ­medicine [96]. Chitosan, poly-(β-1/4)-2-amino-2-deoxy-d-glucopyranose, is a linear polysaccharide composed of d-glucosamine and N-acetyl-d-glucosamine subunits linked by β(1,4) glycosidic bonds [97–99]. The synthesis of chitosan comprises two steps: (1) extraction of chitin (β-[1–4]-poly-N-acetyl-d-glucosamine) followed by the dissolution of calcium carbonate (CaCO3) from crustaceans’ shells with diluted hydrochloric acid, and (2) the deacetylation of chitin through its treatment with diluted aqueous NaOH at 110–115°C for several hours without oxygen. Chitosan is obtained when the degree of deacetylation (DD) is higher than 50% (Fig. 8.6) [100]. Chitosan DD, which can be measured as the amount of free amino groups of the polymer chain, together with the number of protonated amino groups, plays a key role in the solubility, biodegradability, reactivity, and absorption capability of the polymer [99]. Hence chitosan is insoluble in aqueous media above neutral pH, whereas at acidic pH it becomes soluble due to the protonation of free amino acids [81]. In addition, the protonated amino groups can electrostatically interact with negatively charged molecules (e.g., phosphates, sulfates, and citrate ions) or natural polymers (e.g., gelatin, collagen, and fibroin) to form hydrogels [78,99,101]. Chitosan comprises interesting biomedical applications due to its bioactivity, biocompatibility, biodegradability, antibacterial and nonimmunogenic properties, ability for cell ingrowth, and no toxicity. In bone tissue engineering, chitosan has been concretely used as sponges, fibers, films, foams, hydrogels, and particles. It has been shown that chitosan promotes the attachment and proliferation of bone-forming osteoblast cells [78,99,101]. On the contrary, the main drawback of chitosan-based biomaterials is the low mechanical strength hampering its use as a load-bearing structure. As with other natural polymers, the biomineralization of chitosan or its blends with other polymers (e.g., starch) has been used to tackle this limitation [28,102].

8.2.7   Starch Starch is a renewable and cheap natural polysaccharide that can be found in some plants (e.g., corn, potato, wheat, and tapioca) [27,103]. It is synthesized inside plastids CH2OH

CH2OH O

HO

O

O HO

NH

NH O

CH3

Chitin

O

O

HO n

CH2OH

CH2OH

NaOH/heat

O

O

O HO

NH2 O

CH3

O

NH

n

CH3

Chitosan

Figure 8.6  Chemical reaction describing the transformation of chitin into chitosan. Reprinted from Ruiz GAM, Corrales HFZ. Chitosan, chitosan derivatives and their biomedical applications. In: Shalaby EA, editor. Biological activities and application of marine polysaccharides. Rijeka: InTech; 2017. p. Ch. 05 with permission from INTECH.

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through interactions of multiple biosynthetic enzymes and is then deposited as granules into bigger depots of seeds, roots, tubers, fruits, and leaves, among others. Starch has been used in the food industry and as a source of energy since its degradation produces oligosaccharides that can be readily metabolized to obtain energy. Starch is a polymeric carbohydrate composed of a high number of d-glucose units joined together by glycosidic bonds. It contains two types of glucose: α-amylose (20%–30%) and amylopectin (70%–80%). α-Amylose is a linear polymer of several thousands of glucose residues linked by α(1 → 4) bonds, whereas amylopectin is a branched polymer of α-d-glucopyranosyl units composed of α(1 → 4)-linked glucose residues and α(1 → 6) branch points (Fig. 8.7) [104,105]. Starch is low cost, biocompatible, nontoxic, and biodegradable [81,107]. However, starch by itself is brittle and its use in bone tissue engineering presents various drawbacks because of its lack of processability, low surface area, and high water sensitivity [108]. The incorporation of natural polymeric fibers, synthetic polymers, or ceramics has been used to overcome these deficiencies. For instance, its biodegradation, bioactivity, biocompatibility, and mechanical properties can be enhanced by combining starch with vinyl monomers or hydroxyapatite [109]. The presence of starch in mineralized starch–chitosan composites induces an enhancement of compressive strength, swelling ratio, and an increase in carboxyl content, which favors cell proliferation, alkaline phosphate (ALP) activity, and mineralization of osteoblast-like cells [102]. Polycaprolactone blended with starch and loaded with allogeneic goat or rat bone marrow stromal cells has been demonstrated to induce the formation of new bone in large defects [110,111]. Amylose

O OH

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O

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O

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Figure 8.7  Chemical structure of amylose and amylopectin of starch. Reprinted from ref Xie F, Pollet E, Halley PJ, Avérous L. Advanced nano-biocomposites based on starch. In: Ramawat KG, Mérillon J-M, editors. Polysaccharides: bioactivity and biotechnology. Cham: Springer International Publishing; 2015. p. 1467–1553 with permission of Springer International Publishing.

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8.3  Bone regenerative therapies with multifunctional biomaterials of natural polymers In the last few years, the design of synthetic bone grafts shifted from inert materials with high load-bearing capacity (e.g., metallic implants) to bioactive and biodegradable biomaterials that communicate with the body at the cellular level, inducing the regeneration of native tissues and restoring its original functionality [58,112]. In this chapter, we describe the most relevant biomaterial formulations used for bone repair: i.e., scaffolds, hydrogels, and micro-/nanospheres and their combinations, focusing on the use of natural polymers. Subsequently, we focus attention on the principal stimulus proposed as contributors or mediators for improving cell response and bone tissue formation (Fig. 8.8). These strategies have been classified as follow: physical and chemical stimulus (i.e., cross-linking and biomimetic mineralization) to improve the mechanical performance and interfacial properties; and biological strategies to enhance cell activity, which involves loading with bioactive compounds (mainly genes, biomolecules, and growth factor) and cell encapsulation [113].

issue scaffolds ne t o B Hydrogel Nanofibers

P h y sic o m e c h a n i c

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Figure 8.8  Common formulations to obtain biomaterials (i.e., scaffold, hydrogel, and microsphere) for bone repair. The sketch also shows the most commonly used physicomechanical and biological strategies pursued to enhance osteoinductivity of synthetic bone grafts. MMC, macromolecules. Reprinted from ref Fernandez-Yague MA, Abbah SA, McNamara L, Zeugolis DI., Pandit A, Biggs MJ. Biomimetic approaches in bone tissue engineering: integrating biological and physicomechanical strategies. Adv Drug Delivery Rev 2015;84:1–29 with permission of Elsevier.

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8.3.1  Formulations based in natural polymers for bone repair 8.3.1.1  Scaffolds Scaffolds can be defined as a temporary 3D porous structure that provides mechanical support and stimulates cellular colonization and formation of new bone tissue at the same time that it is biodegraded, avoiding the need for further clinical interventions to remove the implant. Scaffold microstructure (i.e., pore size, porosity, permeability) will dictate the surface area available for cell migration from surrounding bone tissues to the inner areas of the scaffolds, which is needed to induce the formation of new bone and its vascularization [114]. The minimum pore size required for new bone formation is 200 μm, while pore interconnections larger than 100 μm are required to ensure suitable permeability for the nutrient and metabolic waste transport and to provide free space for neovascularization [4,18,115]. Freeze drying is the most common methodology used to develop biopolymer-based 3D scaffolds with tailored pore size, porosity, and pore distribution (aligned or isotropic structures). This technique consists of freezing a water-based suspension to induce the nucleation and growth of ice crystals that push aside the solid phase and prompt its concentration between the growing ice crystals. Afterward, the removal of ice via sublimation creates a space generating the porous structure (Fig. 8.9) [116]. Parameters such as freezing temperature, freezing rate, or solute concentrations will strongly affect the nucleation and crystal growth processes, which in turn will influence the final porous scaffold structure. Many examples of collagen scaffolds with tailored pore size and porosity obtained by freeze drying can be found in the literature [115,117]. The fact that collagen is the main organic component of bone ECM makes

Growing lamellar crystals

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Growth direction

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Figure 8.9  Phase diagram of water illustrating the processing steps of freeze casting. Reprinted from ref Deville S. Freeze-casting of porous ceramics: a review of current achievements and issues. Adv Eng Mater 2008;10:155–169 with permission of John Wiley & Sons.

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this biopolymer an ideal candidate to build effective scaffolds [118]. Indeed, many commercially available collagen-based materials can be found in the market (e.g., INFUSE, OP-1 Stryker, Duragen), as previously mentioned in Section 8.2.1. Silk, chitosan, and alginate scaffolds have also been proposed for bone tissue engineering. These polymers have been functionalized (i.e., incorporation of RGD sequence) to improve and promote cell–material interaction [72,83,119–122].

8.3.1.2   Hydrogel Hydrogels are cross-linked 3D networks of hydrophilic polymer chains capable of absorbing large quantities of water or biological fluids while retaining their 3D structure [123]. This 3D structure, similar to the native ECM, makes hydrogels ideal biomaterials for tissue engineering and regenerative medicine. They can be introduced into the body in a minimally invasive manner to fill irregular large defects or to deliver gradually drugs and/or growth factors. On the other hand, the use of hydrogels is still limited because of their low mechanical strength and fragile nature. Hydrogels can be made of synthetic or natural (mainly alginate, gelatin, and chitosan due to their gelling properties) polymers. They can be divided into physical or chemical hydrogels depending on their nature [124]. Physical hydrogels are polymeric networks resulting from the physical cross-linking process (i.e., hydrophobic association, chain aggregation, crystallization, polymer chain complexion, and hydrogen bonding) associated with changes of temperature (i.e., gelatin hydrogels), pH, ionic concentrations, etc. [125] Chemical hydrogels are obtained through enzymatic cross-linking, Schiff base cross-linking, Michael additions, click chemistry, and photo-cross-linking [126,127]. Whereas physical hydrogels are reversible, chemical hydrogels are permanent and irreversible and show better mechanical performance. In the last few decades, stimuli-sensitive hydrogels (referred as smart hydrogels) have attracted special attention in the biomedical field [123]. The 3D structure, mechanical strength, and permeability of smart hydrogels are sensitive to external physical (e.g., light, pressure, temperature, electric or magnetic fields, mechanical stress) and chemical (e.g., pH, ionic factors, and chemical agents) stimuli [128]. Among the stimuli-sensitive hydrogels, the in situ forming hydrogels are of special relevance in the biomedical field since injectable liquid formulations forming a macroscopic gel in situ at the site of injection improve patient comfort and reduce injury cost [128]. For instance, glycerophosphate-based chitosan thermosensitive hydrogels have been proposed as a suitable injectable biomaterial for drug delivery and bone tissue engineering due to the capability of β-glycerophosphate to induce a sol–gel transition in chitosan solutions at physiological pH and temperature [129]. In the case of hydrogels for the delivery of active species, the incorporation of micro- and/or nanoparticles is an efficient strategy to control the release kinetics [126]. For instance, the incorporation of β-tricalcium phosphate (TCP) particles into an alginate gel containing mesenchymal stem cells has been demonstrated to have an osteoinductive effect on soft tissue after subcutaneous implantation in nude mice [127].

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8.3.1.3  Micro-/nanosphere A large variety of nanospheres (with diameters up to 200 nm) and/or microspheres (with diameters ranging from 1 to 1000 μm) have been used for the controlled delivery of drugs or other active species at specific sites or organs [130]. In BTE, the use of micro-/nanospheres has also attracted special interest in the last few decades due to their advantages compared with bulk scaffolds and the possibility of preparing injectable and/or moldable formulations to be applied by using minimally invasive surgery. Their inherently small size and large specific surface area allow the loading of large amounts of active species (see Section 8.3.2.2). Moreover, microspheres are used as cell delivery vehicles since they can encapsulate cells inside and/or attached to their surface (see Section 8.3.2.2). Their use can also increase in vivo reactivity toward the surrounding media, since they can be specifically designed to be sensitive to external stimuli (i.e., magnetic fields, ultrasound, light, electric pulses) or internal signals (i.e., enzyme concentration and/or redox gradients, pH, temperature) [131]. Moreover, micro-/nanospheres can be added to scaffolds and hydrogels to enhance their porosity and/or mechanical properties [130]. Nanoparticles can be obtained through covalent or ionic cross-linking, polyelectrolyte complexation, and self-assembly of hydrophobically modified polymers. Various polysaccharides have been used to produce nanoparticles for drug delivery systems such as dextran, starch, alginate, chitosan, or gelatin [132]. Microspheres can be produced by the emulsification process (i.e., single or double), polymerization, photopolymerization, phase separation/coacervation, spray drying, and solvent extraction [133]. During the production of microspheres, pore interconnectivity, surface topography, surface chemistry, or particle size can be tailored to improve the microspheres’ functionality [133,134]. Microspheres of collagen, gelatin, alginate, and chitosan have been synthesized for biomedical applications [130]. These polymers are not only biocompatible and biodegradable as thoroughly discussed in this chapter, but also contain numerous carboxylic and amine groups, which facilitate the interaction with growth factors or drugs [130].

8.3.2  Modulating the scaffold performance to improve cell activity 8.3.2.1  Physicochemical strategies: cross-linking and bioinspired mineralization Cells are very sensitive to the structural (i.e., morphology, topography, porosity) and compositional (i.e., biomimetic hydroxyapatite containing magnesium, carbonate, or even strontium) features of the surrounding environment. Moreover, cells involved in bone formation and degradation (i.e., osteoclasts, osteoblasts, osteocytes, and mesenchymal stromal cells [MSCs]) are mechanosensitive [135,136]. Therefore special attention has been paid to improving the mechanical performance of synthetic bone substitutes based on natural polymers. The most common chemical strategies such as cross-linking and biomimetic mineralization to efficiently modulate the mechanical response of the biopolymer are thoroughly described herein [5].

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Cross-linking of natural polymers Cross-linking reagents contain reactive ends that specifically react with functional groups of the natural polymers (e.g., primary amines or carboxylic groups) forming bonds between two polymeric chains. The most common chemical cross-linkers are glutaraldehyde [137,138], diisocyanates, carbodiimides (e.g., 1-ethyl-3-[3-dimethyl aminopropyl] carbodiimide and N-hydroxysuccinimide) [139–141], polyepoxy compounds (e.g., 1,4-butanediol diglycidyl ether and phenyl glycidyl ether) [142], acyl azide, and natural cross-linking agents like genipin [143–147]. Although chemical cross-linking usually exhibits high efficiency, their use is currently decreasing due to the formation of residual compounds on degradation in vivo that increases the cytotoxicity of the biomaterial [148,149]. Physical cross-linking consisting of dehydrothermal treatment, ultraviolet and gamma irradiation [150–152], or biological cross-linking with enzymes (e.g., transglutaminase) [153,154] is also widely used to avoid the formation of cytotoxic by-products. In fact, the use of enzymes like transglutaminase has aroused special concern on the development of in situ forming hydrogels [155–157].

Biomimetic mineralization of natural polymers The complex process of bone biomineralization has been proposed to be specifically governed by the action of the collagen matrix, noncollagenous proteins, and bone cells [21,34,158,159]. The resulting organic/inorganic composite tissue exhibits unique biomechanical properties due to the combination of the toughness of the organic collagenous matrix with the compressive strength of the apatite nanoplatelets, preferentially aligned with the c-axes parallel to the longitudinal axis of the fibrils. Natural polymers are very similar in terms of composition to bone ECM proteins. As described in Section 8.2, they are rich in amino acids, which can act as nucleation sites during biomimetic mineralization. The synthesis of materials through routes inspired by in vivo bone mineralization can provide hybrid scaffolds with similar structure and composition of bone at the nanoscale, i.e., with a good integration of the mineral phase into the organic matrix. In fact, biomimetic mineralization of natural polymers has been widely explored not only to enhance the biomechanics of the scaffold, but also to modulate their surface chemistry and topography. These are key factors mediating the cell adhesion of the implant, and thus important features to be finely tuned for triggering the formation of new bone [55,160–162]. It has been demonstrated that the biomimetic mineralization of natural polymers provides nanocomposites with enhanced mechanical properties and similar nanotopography and chemical composition compared to native bone [5]. An interesting biomimetic route consists of dissolving collagen monomers in phosphoric acid solutions and dropping this solution into a calcium hydroxide suspension containing magnesium chloride. Gradual mixing enables the assembly of collagen fibers occurring simultaneously with the precipitation of apatite nanocrystals. The resulting fibrous mineralized composite exhibits compositional (i.e., presence of Mg ions such as in bone apatite) and structural features similar to those of newly formed bone [163–165]. This hybrid scaffold is currently commercialized as RegenOss. It has been safely used to achieve good arthrodesis when associated with autologous bone graft to obtain long spinal fusion in the treatment of adult scoliosis (Fig. 8.10) [166]. The vast majority of

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the commercially available natural polymer–ceramic composites (Table 8.2) are based on the mineralization of collagen type I. Several biomimetic approaches to mineralize matrices of gelatin, chitosan, silk, or alginate have been explored [168,169]. For instance, hydroxyapatite/chitosan scaffolds have been shown to improve the adhesion, spreading, and proliferation of bone marrow-derived mesenchymal stromal cells (BMSCs), inducing their osteogenic differentiation in vitro and in vivo [170]. Biomimetic mineralization of RCPs has also been explored to design bone-like scaffolds. Linear porous scaffolds synthesized via RCP biomimetic mineralization and freeze drying have proved to be osteoconductive and promote osteoblast mineralization in vitro. This study demonstrated that the architectural cues influence cellular proliferation, while the scaffold chemistry and mechanics contribute to gene expression [55]. RCP mineralization in the presence of magnesium turned into hybrid matrices with outstanding interfacial properties (e.g., similar surface roughness to femur bone) favoring the osteogenic differentiation of murine MSCs in vitro [56,57].

Product

Polymer

Ceramic

Recommended use

Collagraft (Zimmer/NeuColl)

Type I (bovine) collagen

HA, TCP

Collapat II (BioMet Inc.) FormaGraft (Maxigen Biotech Inc.) Integra Mozaik (Integra OrthoBiologics) Vitoss (or) Vitoss Bioactive (Orthovita) Mastergraft matrix (Medtronic) CopiOs (Zimmer)

Type I (calf skin) collagen Type I collagen 20% type I collagen

HA HA, TCP 80% TCP

Acute long bone fractures and traumatic osseous defects Aseptic enclosed metaphyseal bone defects Bone void filler Bone void filler

20% collagen

Biostite (Vebas)

Type I (equinine) collagen, chondroitin-6-sulfate

80% β-TCP (or) 70% β-TCP/10% BG BCP Calcium phosphate, dibasic calcium phosphate HA

Bio-Oss Collagen (Geistlich Biomaterials) TricOs T (Baxter CycLos (Mathys Orthopaedics Ltd.) Cerasorb (Curasan Regenerative Medicine) Healos (Depuy Spine) RegenOss (JRI Orthopaedics)

10% (porcine) collagen

HA

Fibrin Sodium hyaluronate Collagen

BCP β-TCP β-TCP

Type I collagen Type I collagen fibers

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Collagen

Nano-HA coating Magnesium-enriched HA nanocrystals Nano-HA

Type I (bovine) collagen Type I (bovine) collagen

Bone void filler, spinal and trauma surgery

Natural polymers for bone repair

Table 8.2  Commercially available natural polymer–ceramic composites

Bone void filler Bone void filler Filling of peridontal defects, preprosthetic osseous reconstruction, maxillofacial reconstructive surgery Filling of periodontal defects, alveolar ridge reconstruction Bone void filler Bone void filler Filling, bridging, reconstruction, and bone fusion Bone void filler, spinal surgery Long bone fractures, revision hip arthroplasty to fill acetabular defects, and spinal fusion Bone void filler

219

BCP, bioactive collagen peptide; BG, bioactive glass; HA, hydroxyapatite; TCP, tricalcium phosphate. Reprinted from reference Yunus Basha R, T.S SK, Doble M. Design of biocomposite materials for bone tissue regeneration. Mater Sci Eng C 2015;57:452–63 with permission of Elsevier.

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8.3.2.2  Biological strategies: delivery of growth factors and cell encapsulation Endogenous signaling molecules are often not enough to repair large bone defects. In such cases, bone regeneration requires the use of biological strategies to deliver on demand exogenous signaling molecules to promote osteoinductivity. There are several devices proposed as delivery systems of biologically active signaling molecules, such as films, particles (micro and nano), hydrogels, and/or scaffolds [171]. Natural polymers have also been used for delivery because of their characteristics such as biocompatibility, biodegradation, high content in charged amino acid groups, and wettability [132]. As discussed in Section 8.3.1.3, the bioactive species can be encapsulated into nano- or microparticles or adsorbed on their surface to be then gradually delivered in spatial, temporal, and dosage-controlled fashions, thus improving efficacy and safety [172]. Several growth factors are involved in bone regeneration, such as BMPs, transforming growth factor, and VEGF. Among them, BMP-2 is well known to play a relevant role in bone healing by influencing osteogenesis and vascularization and leading to new bone tissue formation [131]. The available BMP-2 carriers use supratherapeutic doses and prompt noncontrolled release over time resulting in severe side effects [173]. Scientific interest is therefore focused on innovative systems able to bind, deliver, and guarantee a controlled release of anabolic factors [173]. Leeuwenburgh and coauthors demonstrated the potential of nanostructured colloidal gelatin gels for the sustainable delivery of BMP-2 [66]. In fact, due to its gelling and viscoelastic properties, gelatin hydrogels have been proposed as potential scaffolds for the controlled and localized release of various cytokines (e.g., BMPs, fibroblast growth factors, VEGF) and as cell carriers to treat bone repair and angiogenesis [70]. Bacterial cellulose scaffolds loaded with BMP-2 show great effectiveness in ectopic bone formation and may have potential clinical applications in the treatment of bone defects [174]. In addition to the delivery of growth factors, the seeding of autologous cells into natural polymer platforms has been proposed to accelerate new bone tissue formation. In 1993, the concept of tissue engineering was formally proposed by Langer and Vacanti [112]. It consists of isolating cells from the patient (in particular BMSCs), expanding and seeding them on synthetic scaffolds, and later implanting the cell/ scaffold construction on the bone defect. It has been demonstrated that the delivery of human bone marrow stromal cells (hBMSCs) seeded onto VEGF/BMP-2 releasing composite alginate scaffolds enhances the bone regenerative capability in a criticalsized femur defect [175]. Ben-David et al. demonstrated that a gelatin-based hydrogel scaffold supports the viability and osteogenic differentiation of rat bone marrow-derived cells in vitro and the implantation of this cell–scaffold construction in a critical cranial bone defect enhances new bone formation [176]. Another interesting example is the encapsulation of hBMSCs in chitosan–collagen hydrogels to promote the upregulation of various osteogenic markers such as osterix, bone sialoprotein, and ALP activity [177].

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8.4  Outlook and future perspectives: natural templates fabricated by 3D bioprinting Bone grafts for augmenting or stimulating the formation of new bone in defective areas represents a premier clinical need, as vast numbers of patients are left with insufficient healing of bony defects due to congenital abnormalities, traumatic injury, or bone diseases. As discussed throughout this chapter, conventional strategies for bone repair have found difficulties in synthesizing scaffolds with homogeneous and reproducible porous shape and size distributions and/or tailored surface features. In this regard, additive manufacturing (AM) with a high degree of automation, good accuracy, and reproducibility is an interesting strategy to fabricate bone substitutes with a predefined and customized geometry, size, porous structure, and topography [178–180]. AM consists of building complex 3D structures (made of metals, ceramics, polymers, or composites), usually layer by layer, following a predefined 3D model. 3D bioprinting is an AM-based technology to develop accurate 3D templates and it is currently widely employed to enhance applicability and produce cell-laden scaffolds [181]. The following steps are needed to “bioprint” customized scaffolds: (1) collect the information of fractured or diseased bone by medical imagining techniques (e.g., magnetic resonance, computed tomography scan); (2) design a model of the scaffold with the specific software (i.e., computer-aided design and/or computer-aided manufacturing); (3) transform the 3D model into standard tessellation language, which is then sent to the bioprinter to build the 3D scaffold [80,182]. Bioprinting technology allows the construction of biomimetic 3D structures composed of natural polymers, biologically active molecules, living cells, or any combinations of them, with the possibility of mimicking the biological and compositional features of the host tissues [183–185]. To achieve the best-in-class properties of the final device it is relevant to consider various factors, such as printing fidelity, stability, cross-linking time, biocompatibility, cell encapsulation and proliferation, shear-thinning properties, and mechanical properties (i.e., mechanical strength and elasticity) [181]. On the other hand, an optimal bioprinted scaffold may present similar features to the ECM to sustain cell and tissue growth (i.e., cell adhesion, proliferation, differentiation, tissue growth, and tissue functionalization), appropriate mechanical strength (i.e., sol–gel transition, self-standing 3D structures, loading of growth factors), and adequate mass transfer (i.e., metabolites, nutrition, and growth factors) [79]. Collagen, gelatin, and silk fibroin are broadly used to produce bone substitutes by 3D bioprinting [107]. For instance, cell-laden collagen bioink combined with a nontoxic cross-linking agent (i.e., tannic acid) was proposed to obtain a highly porous structure with improved cell viability and cell distribution [186]. Kim and coauthors fabricated a porous hybrid structure consisting of a layer of cell-laden collagen bioink printed onto an α-TCP/collagen layer without cells. This bioprinted composite scaffold showed good mechanical properties and good cell viability (Fig. 8.11(a)) [187]. On the other hand, Dai and coauthors proposed 3D bioprinting to produce BMSC-laden methacrylamide gelatin scaffolds with collagen-binding domain-bone

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morphogenetic protein-2 (CBD-BMP2) collagen microfibers. BMSC showed high cell viability during the printing process and CBD-BMP2-collagen microfibers promoted BMSC differentiation into osteocytes after 14 days of cell culture (Fig. 8.11(b)) [188]. New advances are required to improve precision during printing, as well as the stability, flexibility, innervation, and vascularization of the “bioprinted” materials [189]. Nevertheless, the development of novel 3D matrices containing cells and biologically active molecules by additive manufacturing is a very interesting route to obtain “personalized” biomaterials. Moreover, this technique presents a more cost-effective form of treatment for patients with musculoskeletal defects or diseases, difficult to treat by conventional therapy [190].

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