CHAPTER TWO
Enzymatically Synthesized Inorganic Polymers as Morphogenetically Active Bone Scaffolds: Application in Regenerative Medicine Xiaohong Wang1,*, Heinz C. Schröder1, Werner E.G. Müller1,*
1ERC Advanced Investigator Grant Research Group at the Institute for Physiological Chemistry, University Medical Center of the Johannes Gutenberg University Mainz, Germany *Corresponding authors: E-mail: [email protected], [email protected]
Contents 1. Introduction29 2. Bone Scaffolds 30 2.1 Bioinert Materials 31 2.2 Bioactive Materials 33 2.3 Regenerative Functional and Custom-Made Tissue Units 35 3. Bone Cells 39 3.1 MSCs/Osteoblasts 39 3.2 Osteoclast Cell Differentiation 41 4. Biogenic, Morphogenetically Active Inorganic Polymers 43 4.1 Biocalcite 43 4.2 Bio-polyphosphate 43 4.3 Biosilica 47 5. Enzymes Controlling the Synthesis of Morphogenetically Active Inorganic Polymers: A Paradigm Shift in Bioinorganic Chemistry 50 5.1 Carbonic Anhydrase 51 5.2 Alkaline Phosphatase 57 6. Biocalcite as Bioseed during Mammalian HA Formation 60 7. CA Activators as Potential Novel Drugs to Stimulate Bone Mineral Formation 62 8. ALP Activators—Potential Novel Compounds to Stimulate Bone Mineral Formation? 63 9. Applications for Bioprinting Organs 65 10. Concluding Remarks 67 Acknowledgments68 References68 International Review of Cell and Molecular Biology, Volume 313 ISSN 1937-6448 http://dx.doi.org/10.1016/B978-0-12-800177-6.00002-5
© 2014 Elsevier Inc. All rights reserved.
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Abstract In recent years a paradigm shift in understanding of human bone formation has occurred that starts to change current concepts in tissue engineering of bone and cartilage. New discoveries revealed that fundamental steps in biomineralization are enzyme driven, not only during hydroxyapatite deposition, but also during initial bioseed formation, involving the transient deposition and subsequent transformation of calcium carbonate to calcium phosphate mineral. The principal enzymes mediating these reactions, carbonic anhydrase and alkaline phosphatase, open novel targets for pharmacological intervention of bone diseases like osteoporosis, by applying compounds acting as potential activators of these enzymes. It is expected that these new findings will give an innovation boost for the development of scaffolds for bone repair and reconstruction, which began with the use of bioinert materials, followed by bioactive materials and now leading to functional regenerative tissue units. These new developments have become possible with the discovery of the morphogenic activity of bioinorganic polymers, biocalcit, bio-polyphosphate and biosilica that are formed by a biogenic, enzymatic mechanism, a driving force along with the development of novel rapid-prototyping three-dimensional (3D) printing methods and bioprinting (3D cell printing) techniques that may allow a fabrication of customized implants for patients suffering in bone diseases in the future.
ABBREVIATIONS 1,25(OH)2D3 1,25-dihydroxy-vitamin D3 3-D Three-dimensional ALP Alkaline phosphatase ASP asialoprotein b-ALP bone-specific alkaline phosphatase BMPs bone morphogenetic proteins BSP bone sialoprotein CA carbonic anhydrase Ca2+ calcium CaCO3 Ca-carbonate CaP calcium phosphate COLI collagen type I CTR calcitonin receptor ECM extracellular matrix GGCX γ-glutamyl carboxylase HA hydroxyapatite HCO3− bicarbonate IκBα nuclear factor of kappa light polypeptide gene enhancer in B-cells inhibitor, alpha M-CSF macrophage–colony-stimulating factor MAPK/ERK mitogen-activated protein kinase/extracellular-signal-regulated kinases MSC mesenchymal stem cells OCAL osteocalcin OPN osteopontin OPG osteoprotegerin
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Pi inorganic phosphate polyP polyphosphates PPi pyrophosphate PPK polyphosphate kinase QA quinolinic acid RANK receptor activator of nuclear factor κB RANKL receptor activator of NF-κB ligand R-Smads receptor-regulated Smads TCP tricalcium phosphate TNAP tissue-nonspecific type alkaline phosphatase TRAP tartrate-resistant acid phosphatase TRIS 2-amino-2-hydroxymethyl-propane-1,3-diol β-GP β-glycerophosphate WNT/SHH WNT/hedgehog
1. INTRODUCTION The ultimate goal for any kind of reconstructive surgery had been tissue/ organ repair from ancient times until present. Repair of tissue/organ defects traditionally involves tissue grafting and/or organ transplantation as well as alloplastic or synthetic material replacement. Since early history until the seventeenth-century gold was used as implant material for hard tissue defects (Sanan and Haines, 1997) but also organically based materials, e.g., the organic marine sponge skeletons/scaffolds, were occasionally applied (Camper, 1771). The manifested limitations of those grafts were tried to compensate later by the application of implants based on synthetic materials of inorganic or organic nature. However, those implants very often failed to integrate into the host tissue and showed inherent disadvantage not to be replaceable by the body’s own cells and tissues. In the 1980s, tissue engineering emerged to overcome those limitations by tissue grafting and/or alloplastic tissue repair (Langer and Vacanti, 1993). In the last two decades, the concept of transplanting of compensatory porous and degradable materials, enriched with biofactors (cells, genes, and/or proteins) has been developed (Hollister, 2005). Attempts that include stem cell approaches and gene therapy approaches followed (Cutroneo, 2003; Audet, 2004; Anam and Davis, 2013).Very recently morphogenetically active scaffolds, suitable for the three-dimensional (3D) growth of mesenchymal stem cells (MSC) and likewise suitable for bioprinting (3D cell printing), have been developed.This chapter outlines strategies to fabricate tissue units by 3D bioprinting technology, as well as a novel approach to apply activators for the recently elucidated major enzymes involved in the biosynthesis of the bone biomineral, the carbonic anhydrase (CA) and the alkaline phosphatase (ALP).
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2. BONE SCAFFOLDS It is the task of tissue engineering to develop biologically active organ/tissue substitutes that have the property to restore lost morphological and functional features of impaired or diseased organs. As outlined by Langer and Vacanti (1993) in tissue engineering new developments of functional substitutes for damaged tissue can only be successfully translated into practice if the basic principles of biology and engineering can be amalgamated to the invention. One discipline alone cannot achieve this goal. This interdisciplinary field should provide solutions for tissue creation and repair. From cell biology it is known that cells composing tissues, from the basis of the metazoan kingdom the sponges (phylum: Porifera) to the crown Metazoa, the mammals, and the insects, are not loosely embedded in the tissue but integrated to functional units by controlled and directed cell–cell interactions as well as cell–matrix interactions (Müller et al., 2004). This biological basic construction network for Metazoa allows intracellular as well as extracellular signaling information transmission networks but also via extracellular matrix (ECM) elements, e.g., collagen and fibronectin, concerted circuits that are regulated by biological, physical, and chemical cues of the microenvironment. Those tuned interactions provide the critical platform for integrated cell functions and behaviors. In turn, scaffolding materials to be designed and to be intended for tissue engineering applications must mimic those physiologic environments. Even more, these circuits not only integrate the cells between soft tissue and hard tissue but also, and there especially, determine the 3D geometrical, topographical, as well as physical units. Physiological matrices as well as fabricated 3D scaffolds are the crucial prerequisites to elicit, induce, and trigger the cells with the physiologically relevant stimuli in order to establish, maintain, and further develop their functionalities to associate and to build tissue. In turn, a shift of paradigm is presently proceeding which builds on (1) the collected experiences from the established bioinert scaffold materials as well as (2) the knowledge on bioactive materials and is culminating in (3) the development of regenerative and custom-made biosynthetic implants and tissue grafts. It is the aim to fabricate scaffolds for engineered tissue units using biodegradable soft and simultaneously porous materials that allow the embedding and integration of biological cells with growth and differentiation factors, exogenously added or synthesized by the cells themselves.
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2.1 Bioinert Materials The application of bioinert materials for tissue and bone tissue replacement is old. Especially the application of biomaterials from marine animals as plaster and in bone/tissue replacement has been a tradition since the Greek times. Camper (1771) described that the organic matrix of sponges can be successfully used in plastic surgery of the palate of the skull (Figure 2.1(a)). He was fabricating a nose using lime wood, covered it with a sponge and fixed it in the roof of the mouth via a silk thread which had been waded by small sponge slices. It might also be noted that the application of the siliceous skeleton of sponges as suitable scaffold onto which human stem cells can be seeded has recently been reported (Green et al., 2003). Basically two basis materials have been used to fabricate implants for bone reconstitution/reconstruction; (1) metals and (2) ceramics. In the first phase of those replacement supports, the intention for the use of those implants had been to stabilize the body at the position of the damaged bone (Chang et al., 1996). For decades metal implants have been used in orthopedics for mechanical skeletal repair. Those supports had to meet the challenge to strengthen the implant–bone interface and to prevent stressshielding effects. Those implants can be fabricated in customized processes, e.g., by 3D printing (Figure 2.1(b)). The key issue for a durable and successful implant is the establishment of a strong bone–implant interface. It emerged that smooth implant surfaces can result in the formation of encapsulation with the consequence of loosening of the implant (Greco, 1994). In one approach to extend and to promote long-term interface strength, porous materials and porous coatings have been developed (Engh and Bobyn, 1985). Those porous materials and coatings induce a partial to complete bone ingrowth, which has the advantageous property to enhance the strength of the interface bonding under simultaneous reduced tendency to cause capsule formation around the implant. A further challenging issue for most of those metalbased implants is the appearance of “stress shielding” (Jacobs et al., 1993; Amstutz, 1991). Even though it is well established that bone regeneration and repair processes are promoted by mechanical loads (Van Lenthe et al., 1997) metal materials such as titanium, still widely used today for bone implants, is much stiffer than native bone. Consequently, an implant of solid titanium can carry a disproportionate amount of the biological loads (see: Thelen et al., 2004). In turn, the surrounding bone undergoes a process of “stress shielding” and suffers from abnormally low levels of
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A
B
C
D
Figure 2.1 The development of bone scaffolds. (a) Bioinert materials, e.g., application of sponges in tissue replacement (Camper 1771). The damaged nose/nostrils (A, B, D, C) was modeled by a piece of lime wood covered with a sponge (T, U, V), and fixed in the roof of the mouth (W) via a silk thread (S) that had been surrounded with small sponge slices. (b) Computer-aided rapid prototyping/3D printing. (A) Data are generated for an organ or tissue unit using the computing process. Algorithms for the automated design and fabrication of a scaffold/bone part are developed based on an assembly-free process. (B to D) The bone unit is printed, like in an ink-jet printer, and comprises a tightfitting customized morphology. If it is formed of ceramic or titanium, the implant has usually only an osteoconductive property. (c) Fabrication of an osteoinductive scaffold made of glass, metal, or ceramic that allows the cells to migrate into its pores. The scaffold is coated with bioactive factors or polymers, e.g., bone morphogenetic proteins-2 or polyphosphate, that direct its associated cells, stem cells, to terminally differentiated cells that build tissue units like blood vessels.
stress, which finally results again in bone resorption, followed by loosening of the implant (Black, 1999). A further achievement toward a more advanced material to be used for orthopedics is the use of bioinert ceramics (Boutin, 1981). Ceramic, an
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inorganic, nonmetallic solid that is prepared by the action of heat and subsequent cooling, has been used since the 1970s in orthopedics (Smith, 1963; Eyring and Campbell, 1969). As material, ceramics could have a crystalline or partly crystalline structure, or be amorphous (e.g., a glass). Ceramics are increasingly used in the orthopedic surgery, especially for joint prostheses. It is used not only for joint bearings, but also for the bone–implant interface of prostheses (see: Hayashi et al., 1993), especially for implants as knee prosthesis, total ankle prosthesis, or total elbow prosthesis. The bioinert ceramics have been found to have the property of excellent resistance during carrying (Kumar et al., 1991). Since the beginning of the development of ceramic implants it is hoped that this material is more biocompatible than metal alloys since it is provided with the property of resistance to corrosion, to be less cytotoxic and to be hydrophilic. At present synthetic biodegradable polymers, interconnected with porous calcium hydroxyapatite (HA) ceramics have been found to be very suitable composite materials for implants, since they can be combined with growth and development factors in a carrier/scaffold system, e.g., with recombinant human bone morphogenetic protein-2 (rhBMP-2), that strongly promotes the clinical effects of rhBMP-2 in bone tissue regeneration (Bianco and Gehron, 2000). This study, like others that appeared during this time period (Hollister, 2005), predetermines the development to a second level of regenerative restorative implantology the “bioactive materials.”
2.2 Bioactive Materials The strategy to design more effective bioactive tissue engineering scaffolds is to implement in a complex manner three essential elements, first, a porous matrix (scaffold), second, to elicit osteoconductive signals, and third, to implant, adjacent to the matrix, osteogenic cells that can attach to the matrix and respond to their signals via an adequate blood supply (Finkemeier, 2002); Figure 2.1(c). In turn, the implants must be designed in a hierarchical way; first hierarchical porous structures must be prepared that are provided with a suitable mechanical stability and flexibility and allow the transfer and diffusion of growth factors and differentiation factors. The complex 3D anatomical shape of the bone substitution material must try to imitate from the nanometer to the millimeter level the functional properties of the natural bone.The scaffold formed must ideally meet the requirements of the cells to be provided with nutrients. The porous channels must allow cell migration, and their surface features must be suitable for cell attachment (Cukierman et al., 2001). The morphology/topography as well as the
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roughness of the scaffold surface should fall between the theoretical maximum given by the material and the theoretical minimum of zero predicted by composite theories (Torquato, 2002; Hollister, 2005). Consequently, the critical issue for the design of the surface texture is based on computer calculations that match the requirements to allow intercellular and transcellular signal transmission as well as to leave space for the development of a vascular system that allows mass transport. An effective permeability is determined by the 3D pore arrangement and the adhesion receptors associated with the plasma membrane, e.g., the integrins. The biological properties of the surfaces of the, basically inert, matrices are crucially important.The material must be bioactive along an increasing complexity and inducibility, from being osteoconductive to osteoinductive and allowing processes that cause osseointegration. According to the definition by Albrektsson and Johansson (2001) the term osteoconduction means that bone grows on a surface, e.g., bone surface, that supports the ingrowth of the osteoblasts into pores, channels, or pipes (Wilson-Hench, 1987). However, more often the surface used is not the bone but another biogenic growth platform (Glantz, 1987). The next higher level of bioactivity is osteoinduction. This term stands for the activity of a contact or soluble material that displays the potency to induce the undifferentiated and pluripotent stem cells to enter the differentiation pathway toward the bone-forming cell lineage, more specific to induce osteogenesis. Finally, osseointegration refers to the process by which the implant is stably anchored into the bone. At present and strictly speaking, only the bone morphogenetic proteins (BMPs) being members of the transforming growth factor-β (TGF-β) superfamily of growth factors and well established physiological regulators of osteoblastic differentiation (Lavery et al., 2008) can be considered as osteoinductive. Among them BMP-7 and BMP-2 display the highest osteoinductivity (e.g., Urist, 1965; Lavery et al., 2009; Sampath et al., 1990). Already in limited clinical use is BMP-2 that has been shown to induce new bone formation in spine fusions and long bone nonunion fractures (Gautschi et al., 2007). After binding of BMP to the integrated cell surface receptor, a tetramer of serine/threonine kinase transmembrane receptors consisting of two type I and two type II receptors, intracellular signaling occurs via intracellular signaling proteins to the receptor-regulated Smads (R-Smads). In turn R-Smads form heteromeric complexes with the common mediator Smad, Smad-4, and subsequently translocate to the nucleus where they act as transcription factors to induce the BMP responsive genes (Sebald et al., 2004). This implies that any kind of scaffold supplemented/coated with BMPs must be qualified to be
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an osteoinductively acting implant. However, recently it becomes overt that in osteoblasts also BMP-independent anabolically acting routes exist. Examples are polyphosphate (polyP) and biosilica-mediated pathways (Wang et al., 2014a,b), natural inorganic polymers that have the capacity to induce BMP-2 in osteoblasts. In addition, a BMP-independent route has been proposed that induces the differentiation of osteoblast precursor cells to mature functionally active osteoblasts (Müller et al., 2013a). It also appears likely that the phytoestrogen isoquercitrin acts synergistically with polyP, on the transcription factor RUNX2 (Wang et al., 2014c). In line with a broader interpretation of osteoinductivity (Amini et al., 2012) biomaterials, including natural and synthetic ceramics (i.e., HA and various calcium phosphate (CaP) compositions, and their composites), have been qualified as osteoinductive materials. Besides of synthetic CaP-based biomaterials, also in the form of sintered ceramics (Yamasaki and Sakai, 1992; Klein et al., 1994), cements (Habibovic et al., 2008), and coatings (Habibovic et al., 2004), natural and coral-derived ceramics (Ripamonti, 1991; Ripamonti et al., 2009) have been attributed to be osteoinductive. Especially highlighted should be porous bioglass that has been widely introduced into clinics (Jones, 2013). This bioactive ceramics, a biodegradable glass of a general formula “Na2O–CaO–SiO2–P2O5,” contains high levels of calcium (Ca2+); its most generally used formulation contains 46.1 mol% SiO2, 24.4 mol% Na2O, 26.9 mol% CaO, and 2.6 mol% P2O5 and has been termed 45S5 Bioglass (Hench et al., 1971). The bioactive glasses are reported to stimulate bone regeneration to a larger extent than other bioactive ceramics. This 45S5 Bioglass forms a semichemical bond with bone and, in vivo, bonds to other bioceramics. In comparison, the CaP-based materials are more widely used in the clinics.
2.3 Regenerative Functional and Custom-Made Tissue Units Already a decade ago it has been prognosticated that tissue engineering technology, based on computer-aided jet-based 3D organ/tissuelike printing, could be a solution of the organ transplantation crisis (Mironov et al., 2003). However, it is needed to distinguish between inert implants that cannot be replaced by cells/tissue of the recipient and functional graft substitutes that are disintegrated and subsequently replaced by cells and ECM filaments ingrowth from the surrounding tissue of the recipient. Many types of CaP biomaterials have been developed that comprise a similar composition like native bone mineral and its precursors, e.g., HA and α- and β-tricalcium phosphate in the form of ceramics, cements, and thin
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coatings (LeGeros, 2002). A few of these insoluble CaP materials are even osteoconductive and in some cases provided with the ability to induce bone formation, implying that those materials are osteoinductive (Damien and Parsons, 1991; Habibovic and de Groot, 2007). The implant, even though acting to a limited extent osteoinductively, remains as a major core in the previously damaged region.The process of biodegradation around the bone graft substitutes is favored since they prevent the disadvantageous resorption of the neighboring bone due to stress-shielding effects. The technique of free-form fabrication applicable for producing of 3D synthetic bone graft substitutes allows a precise control of the overall geometry and in turn also of the porous structure of the scaffold/implant. The 3D printing of CaP-based structures at room temperature has been successfully demonstrated (Gbureck et al., 2007). By modification of the surface of the implant, either with respect to the morphological, ultrastructural, or chemical properties, the capacity of the material to induce ectopic bone formation can be improved. The fabrication of 3D implants by direct cell printing based on computer-aided design files offers a sophisticated and challenging direction to engineer 3D tissuelike units, to be placed into living human organs (Figure 2.2). However, those approaches require three sequential hierarchical steps of increasing complexity. First, compilation of preprocessing followed by the development of computational “blueprints” of a given organ; second, processing or the program file for the printing of the actual organlike unit; and finally, postprocessing or organ conditioning and ultimately organ maturation. In recent years several types of cell printers have been developed (Tasoglu and Demirci, 2013). The basic principle is that the cells, either in suspension or as aggregates, are embedded into a printable matrix which is then sequentially layered under formation of predesigned blocks. It is surely feasible to fabricate implantable, by rapid-prototyping, 3D organlike units in the future; but it remains open when this concept of tissue engineering can be exploited and integrated into the constraints emerging from the biological, genetic rules of developmental biology. It is just extraordinarily difficult to tailor a suitable matrix into which the cells can be embedded to become provided with the required physiological solute and fibrous extracellular molecules. Only after deciphering the genetic blueprint of the cells and its time- and space-specific expression it will become possible to lay the ground with the chemical and physical cues for the stem/precursor cells to direct them toward an integrative assembly in an organ.
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Figure 2.2 Regenerative functional and customized tissue units bioprinted by the 3D cell printing approach. Sketch of the bioprinting of cells, embedded into a matrix, e.g., alginate/gelatin, using a 3D bioplotter. A cell suspension is filled into a cartridge hooked to the printing head (a). This control element is connected with the computer-guided printing apparatus; the alginate/gelatin/cells are passed through a needle into a CaCl2 bath, which hardens the scaffold (b, c). This scaffold with the bioprinted cells is submersed into medium/serum. Then the 3D scaffold is overlayed with an agarose layer containing the morphogenetically active factors or polymers, e.g., polyP (d). In such an environment the cells proliferate and differentiate.
For 3D bioprinting a critical size of the aggregates has to be intended. A cell density within the organic matrix of >106 cells/ml, according to our experiences, is preferable (Neufurth et al., 2014). The size of the aggregates formed within the scaffold matrix is determined by the supply of nutrients and growth factors and/or morphogenetically active polymers. Usually the aggregates are spherical and can reach sizes of a diameter of 500 μm. Incubation conditions must be developed that are favorable for the cells to spread onto a cell substrate, e.g., fibronectin or collagen. In turn, cell–cell and cell-substrate adhesion must be finely tuned; the cells must express the property to decrease substratum adhesivity while simultaneously allow increasing cell–cell cohesivity, and vice versa (Ryan et al., 2001). In addition, the cell environment should be tailored in a way that one cell type is directed towards cell–cell cohesion while the other cell type undergoes intensive cell spreading. The cell aggregates formed after bioprinting can be composed homocellularly, comprising only one type of cell or heterocellularly, being a hybrid composed of more than one cell type (Figure 2.3(a) and (b)). If homocellular aggregates are formed in a bioprinted matrix the subsequently formed
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Figure 2.3 Bioprinting of cells, embedded in an organic matrix. (a) Either homocellular aggregates are allowed to be formed that have the potency to form basic but simple building units. Alternatively, (b), heterocellular aggregates are allowed to be formed that can differentiate to functional units, e.g., comprising vascular structures (v), within cells derived from a different stem cell lineage. (c) It is still an ambition to bioprint cells of different stem cell origin that can functionally differentiate to different cell types that build an organ.
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tissue structures can show different morphologies but hardly form functional organ units (Figure 2.3(a)). In contrast, if heterocellular aggregates are allowed to be formed within the matrix the formation of internal structures can be expected, e.g., blood vessel within endothelial and smooth muscle cells. Alternatively, peripheral blood stem cells can be used that can differentiate either into endothelial cells or smooth muscle cells; endothelial progenitor cells are present in peripheral blood stem that can be induced to cells supporting neoendothelialization (Sugaya et al., 2012). At present it seems to be manageable to sequentially bioprint cell types with a different phenotype (Figure 2.3(b)). The solution of choice in the future will be to bioprint cells that, based on their cell adhesion properties and origin from a given stem cell lineage, proliferate and in parallel differentiate to a complex organ (Figure 2.3(c)).
3. BONE CELLS Bone appears to be a solid, rigid organ; however, it is highly flexible and dynamic allowing bone anabolic and bone catabolic processes to proceed in a tuned interacting manner. The bone is under continuous remodeling that takes place throughout the life span of an individual. It should be highlighted that the morphology and the shape or the size of bone is genetically determined. Until now it remains unclear which genetic blueprint controls the form-giving processes in bone. Under physiological conditions the net balance between osteoblastic bone formation, mediated by osteoblasts, and osteoclastic bone resorption, driven by osteoclasts, is very much tuned. The bone anabolic cells, the osteoblasts, originate from MSCs having the potential to proliferate and the capacity to differentiate into several connective tissue/cell types. In contrast, the bone catabolic cells, the osteoclasts arise from hematopoietic stem cells (Teitelbaum, 2006).
3.1 MSCs/Osteoblasts The pluripotent MSCs have the potency to differentiate into osteoblasts, chondroblasts, bone marrow stromal cells, fibroblasts, muscle cells, or adipocytes depending on the presence of the growth and differentiation factors in their microenvironment (Wang et al., 2014a); Figure 2.4. Osteoblasts having a cuboidal or columnar shape are lining the bone surfaces at those sites that undergo active bone formation during bone development or fracture repair. Osteoblasts express high levels of type I collagen (COLI) and proteoglycans (glycosaminoglycans), the two main components of the bone matrix, also
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Figure 2.4 Multipotent differentiation of multipotent human mesenchymal stem cells (hMSC). Specific signaling molecules and growth factors as well as differentiation factors induce/activate transcription factors and by that determine both the commitment and the differentiation of hMSCs toward the osteogenic, chondrogenic, adipogenic, or myogenic lineage. The osteogenic and the chondrogenic lineages are involved in the restorative repair of bone and cartilage tissue (osteochondral tissue reconstitution). Biosilica and polyphosphate (polyP) display anabolic, morphogenetic effects on those two differentiation lines. BMP-2, bone morphogenetic proteins-2; ALP, alkaline phosphatase.
termed osteoid. Osteoblasts are also involved in mineralization of osteoid, very likely via the liberation of matrix vesicles, and by the deposition of calcium, carbonate, and phosphate (Landis et al., 1993; Hohling et al., 1978; Müller et al., 2013b). Osteoblasts are aligned by adherens-type junctions, including desmosomes and tight junctions (Safadi et al., 2009). Osteoblasts synthesize and secrete a variety of cytokines and colony-stimulating factors controlling myelopoiesis, e.g., interleukin-6, interleukin-11, granulocyte–macrophage colony-stimulating factor and macrophage–colony-stimulating factor
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(M-CSF). In addition, osteoblasts synthesize a series of growth factors, including TGF-β, BMPs, platelet-derived growth factors, and insulin-like growth factors. Finally, terminally differentiated osteoblasts possess receptors for the parathyroid hormone as well as for 1,25-dihydroxy-vitamin D (1,25(OH)2D3), the major hormones regulating bone metabolism and mineral deposition (Figure 2.5).
3.2 Osteoclast Cell Differentiation Osteoclasts, originating from the hematopoietic lineage (Boyle et al., 2003), undergo differentiation and maturation in the presence of the M-CSF and of the receptor activator of NF-κB ligand (RANKL). Markers for the multinucleated osteoclasts are the highly expressed tartrate-resistant acid phosphatase (TRAP), as well as calcitonin receptor and integrin avb3 (Cerri et al., 2003); Figure 2.5. The cytokine/receptor triad, RANKL with its receptor (RANK) and the endogenous decoy receptor osteoprotegerin (OPG) crucially control bone formation and bone remodeling (Boyce and Xing, 2008; Santini et al., 2011). While RANKL is synthesized by the osteoblastic lineage cells this signaling molecule is essential for the differentiation of those cells that are involved in bone resorption, the osteoclasts. RANKL is expressed on osteoblasts, T cells, dendritic cells, as well as their precursors from where it can be released by specific proteases (Zhang et al., 2009).This ligand (RANKL) binds to the cell surface receptor RANK, located on precursor and mature osteoclasts and by that promotes osteoclastogenesis. After binding of RANKL to RANK the osteoclasts become activated and resorb bone mineral; during this process the cells have close contact to the bone surface (Fuller et al., 2010). At this interphase, osteoclasts to bone, vesicles are formed, via integrin (avb3), that contain proton pumps and acid hydrolases, e.g., cathepsin K. Those enzymes and vesicles are present in those cells that are bone-apposed. “Resorptive hemivacuoles” are formed between osteoclasts and bone, allowing the protons to dissolve the HA scaffold of the bone (Figure 2.5). The intracellular pH is kept close to neutral via the chloride/bicarbonate exchanger.The function of RANKL is under control of OPG, a decoy receptor that is secreted by stromal cells and also by osteoblasts (Kearns et al., 2008). OPG scavenges RANKL by binding to it and neutralizes its function. In turn it has to be concluded that any deregulation of the tuned expression of the RANKL/RANK/OPG system causes a dysregulation of the differentiation pathways of the osteoblasts and the osteoclasts and in turn promotes catabolic bone remodeling (Boyce and Xing, 2008). By that, OPG prevents bone matrix from excessive
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Figure 2.5 Differentiation of the progenitor cells of the bone-forming osteoblasts (osteoblastogenesis) and the bone-resorbing osteoclasts (osteoclastogenesis). Upper panel: Osteoblast differentiation starts from the mesenchymal stem cells and ends with the osteocytes. The major transcription factor Runx2, which is under the control of bone morphogenetic proteins-2, is synthesized in chondrocytes and causes a stage- dependent increase in the structural and functional proteins in osteoblasts, for example, b-ALP (bone-specific alkaline phosphatase), COLI (collagen type I), OP (osteopontin), ASP (asialoprotein), BSP (bone sialoprotein), and OCAL (osteocalcin), as well as RANKL (receptor activator of NF-κB ligand). Lower panel: Principle differentiation stages from the hematopoietic stem cells via preosteoclasts to functionally active, bone-resorbing osteoclasts. The osteoblasts direct the preosteoclasts to the osteoclast through the interaction of RANKL with RANK (receptor activator of nuclear factor κB), an interaction that is blocked by OPG (osteoprotegerin). Differentiation from hematopoietic stem cells starts via activation of the PU.1 transcription factor and inflammatory signals. The CD34+ osteoclast precursor cells, after entering the circulating system and in the presence of M-CSF (macrophage–colony-stimulating factor) and 1,25-dihydroxy-vitamin D3 (vitamin D3), become recruited onto the surface of bone. The preosteoclasts, after the stimulation of the DAP12 adapter protein/receptor undergo multinucleation to the osteoclasts. Those cells express in the presence of 1,25-dihydroxy-vitamin D3 the receptor RANK. After binding of RANKL to RANK the osteoclasts dissolve HA by lowering the pH. Markers for the activated osteoclasts are TRAP (tartrate-resistant acid phosphatase) and CTR (calcitonin receptor). HA, hydroxyapatite.
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resorption by binding to RANKL and in turn abolishes the activation of the osteoclasts via inhibition of the RANK pathway. In conclusion, the relative concentrations of RANKL and OPG in bone are the major morphogenetic determinants of bone mass and strength.
4. BIOGENIC, MORPHOGENETICALLY ACTIVE INORGANIC POLYMERS 4.1 Biocalcite Exemplarily the beneficial function of the calcareous corals in bone reconstruction has been demonstrated (Cooper et al., 2014). Especially the effects of secondary metabolites from soft corals acting against inflammation and tumor growth have been highlighted (Chen et al., 2013). Well understood is the inductive osteogenic differentiation effect of coral scaffold on MSCs (Puvaneswary et al., 2013). Even, with respect to some biological markers, calcareous scaffolds, derived from corals, have been found to be superior in comparison to bone grafts. A significantly higher level of expression of the osteogenic differentiation markers, ALP, osteocalcin (OCAL), and osteonectin, as well as of the transcription factor Runx2 has been described. Even more, the extent of mineralization within coral grafts has been found to be more extensive compared to bone grafts. Further studies revealed that coral products have a curative potential on bone deficits as well. The naturally occurring calcium, within the calcareous scaffold, in the form of aragonite found in the scleractinian hard corals and in the form of calcite deposits within the soft octocorals, contribute to anabolic bone restoration. This effect is especially pronounced if these minerals are administered together with zeolite, a microporous mineral, in mice induced to a menopausal state (Banu et al., 2012); this effect has been confirmed in rabbits as well (Parizi et al., 2012). Parallel with the effect of coral minerals on bone formation, their effect on dental deformities has been studied (Figueiredo et al., 2010). The data indicated that the coral skeleton in its unrefined form cannot be applied due to its necrotic potential. However, if purified coral minerals are applied a beneficial osteogenic effect on bone marrow stromal cells is seen; an effective repair of mandibular defects in canines has been reported (Yuan et al., 2010).
4.2 Bio-polyphosphate Like the chemically synthesized inorganic polymeric phosphate, polyP, the biogenic polyphosphate (bio-polyP) has an amorphous state (Kulaev et al.,
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2004; Rao et al., 2009; Omelon and Grynpas, 2008). In contrast to the chemically synthesized polyP, which is synthesized at high temperature, the biogenically produced bio-polyP is synthesized at ambient physiological conditions (Rao et al., 2009) via polyP kinases; Figure 2.6(a). The biopolymer bio-polyP is found in a wide range of organisms, including bacteria, fungi, algae, plants, and animals (see: Rao et al., 2009); it is readily water soluble in millimolar concentrations at chain lengths
Enzymatically Synthesized Inorganic Polymers as Morphogenetically Active Bone Scaffolds: Application in Regenerative Medicine Xiaohong Wang1,*, Heinz C. Schröder1, Werner E.G. Müller1,*
1ERC Advanced Investigator Grant Research Group at the Institute for Physiological Chemistry, University Medical Center of the Johannes Gutenberg University Mainz, Germany *Corresponding authors: E-mail: [email protected], [email protected]
Contents 1. Introduction29 2. Bone Scaffolds 30 2.1 Bioinert Materials 31 2.2 Bioactive Materials 33 2.3 Regenerative Functional and Custom-Made Tissue Units 35 3. Bone Cells 39 3.1 MSCs/Osteoblasts 39 3.2 Osteoclast Cell Differentiation 41 4. Biogenic, Morphogenetically Active Inorganic Polymers 43 4.1 Biocalcite 43 4.2 Bio-polyphosphate 43 4.3 Biosilica 47 5. Enzymes Controlling the Synthesis of Morphogenetically Active Inorganic Polymers: A Paradigm Shift in Bioinorganic Chemistry 50 5.1 Carbonic Anhydrase 51 5.2 Alkaline Phosphatase 57 6. Biocalcite as Bioseed during Mammalian HA Formation 60 7. CA Activators as Potential Novel Drugs to Stimulate Bone Mineral Formation 62 8. ALP Activators—Potential Novel Compounds to Stimulate Bone Mineral Formation? 63 9. Applications for Bioprinting Organs 65 10. Concluding Remarks 67 Acknowledgments68 References68 International Review of Cell and Molecular Biology, Volume 313 ISSN 1937-6448 http://dx.doi.org/10.1016/B978-0-12-800177-6.00002-5
© 2014 Elsevier Inc. All rights reserved.
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Abstract In recent years a paradigm shift in understanding of human bone formation has occurred that starts to change current concepts in tissue engineering of bone and cartilage. New discoveries revealed that fundamental steps in biomineralization are enzyme driven, not only during hydroxyapatite deposition, but also during initial bioseed formation, involving the transient deposition and subsequent transformation of calcium carbonate to calcium phosphate mineral. The principal enzymes mediating these reactions, carbonic anhydrase and alkaline phosphatase, open novel targets for pharmacological intervention of bone diseases like osteoporosis, by applying compounds acting as potential activators of these enzymes. It is expected that these new findings will give an innovation boost for the development of scaffolds for bone repair and reconstruction, which began with the use of bioinert materials, followed by bioactive materials and now leading to functional regenerative tissue units. These new developments have become possible with the discovery of the morphogenic activity of bioinorganic polymers, biocalcit, bio-polyphosphate and biosilica that are formed by a biogenic, enzymatic mechanism, a driving force along with the development of novel rapid-prototyping three-dimensional (3D) printing methods and bioprinting (3D cell printing) techniques that may allow a fabrication of customized implants for patients suffering in bone diseases in the future.
ABBREVIATIONS 1,25(OH)2D3 1,25-dihydroxy-vitamin D3 3-D Three-dimensional ALP Alkaline phosphatase ASP asialoprotein b-ALP bone-specific alkaline phosphatase BMPs bone morphogenetic proteins BSP bone sialoprotein CA carbonic anhydrase Ca2+ calcium CaCO3 Ca-carbonate CaP calcium phosphate COLI collagen type I CTR calcitonin receptor ECM extracellular matrix GGCX γ-glutamyl carboxylase HA hydroxyapatite HCO3− bicarbonate IκBα nuclear factor of kappa light polypeptide gene enhancer in B-cells inhibitor, alpha M-CSF macrophage–colony-stimulating factor MAPK/ERK mitogen-activated protein kinase/extracellular-signal-regulated kinases MSC mesenchymal stem cells OCAL osteocalcin OPN osteopontin OPG osteoprotegerin
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Pi inorganic phosphate polyP polyphosphates PPi pyrophosphate PPK polyphosphate kinase QA quinolinic acid RANK receptor activator of nuclear factor κB RANKL receptor activator of NF-κB ligand R-Smads receptor-regulated Smads TCP tricalcium phosphate TNAP tissue-nonspecific type alkaline phosphatase TRAP tartrate-resistant acid phosphatase TRIS 2-amino-2-hydroxymethyl-propane-1,3-diol β-GP β-glycerophosphate WNT/SHH WNT/hedgehog
1. INTRODUCTION The ultimate goal for any kind of reconstructive surgery had been tissue/ organ repair from ancient times until present. Repair of tissue/organ defects traditionally involves tissue grafting and/or organ transplantation as well as alloplastic or synthetic material replacement. Since early history until the seventeenth-century gold was used as implant material for hard tissue defects (Sanan and Haines, 1997) but also organically based materials, e.g., the organic marine sponge skeletons/scaffolds, were occasionally applied (Camper, 1771). The manifested limitations of those grafts were tried to compensate later by the application of implants based on synthetic materials of inorganic or organic nature. However, those implants very often failed to integrate into the host tissue and showed inherent disadvantage not to be replaceable by the body’s own cells and tissues. In the 1980s, tissue engineering emerged to overcome those limitations by tissue grafting and/or alloplastic tissue repair (Langer and Vacanti, 1993). In the last two decades, the concept of transplanting of compensatory porous and degradable materials, enriched with biofactors (cells, genes, and/or proteins) has been developed (Hollister, 2005). Attempts that include stem cell approaches and gene therapy approaches followed (Cutroneo, 2003; Audet, 2004; Anam and Davis, 2013).Very recently morphogenetically active scaffolds, suitable for the three-dimensional (3D) growth of mesenchymal stem cells (MSC) and likewise suitable for bioprinting (3D cell printing), have been developed.This chapter outlines strategies to fabricate tissue units by 3D bioprinting technology, as well as a novel approach to apply activators for the recently elucidated major enzymes involved in the biosynthesis of the bone biomineral, the carbonic anhydrase (CA) and the alkaline phosphatase (ALP).
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2. BONE SCAFFOLDS It is the task of tissue engineering to develop biologically active organ/tissue substitutes that have the property to restore lost morphological and functional features of impaired or diseased organs. As outlined by Langer and Vacanti (1993) in tissue engineering new developments of functional substitutes for damaged tissue can only be successfully translated into practice if the basic principles of biology and engineering can be amalgamated to the invention. One discipline alone cannot achieve this goal. This interdisciplinary field should provide solutions for tissue creation and repair. From cell biology it is known that cells composing tissues, from the basis of the metazoan kingdom the sponges (phylum: Porifera) to the crown Metazoa, the mammals, and the insects, are not loosely embedded in the tissue but integrated to functional units by controlled and directed cell–cell interactions as well as cell–matrix interactions (Müller et al., 2004). This biological basic construction network for Metazoa allows intracellular as well as extracellular signaling information transmission networks but also via extracellular matrix (ECM) elements, e.g., collagen and fibronectin, concerted circuits that are regulated by biological, physical, and chemical cues of the microenvironment. Those tuned interactions provide the critical platform for integrated cell functions and behaviors. In turn, scaffolding materials to be designed and to be intended for tissue engineering applications must mimic those physiologic environments. Even more, these circuits not only integrate the cells between soft tissue and hard tissue but also, and there especially, determine the 3D geometrical, topographical, as well as physical units. Physiological matrices as well as fabricated 3D scaffolds are the crucial prerequisites to elicit, induce, and trigger the cells with the physiologically relevant stimuli in order to establish, maintain, and further develop their functionalities to associate and to build tissue. In turn, a shift of paradigm is presently proceeding which builds on (1) the collected experiences from the established bioinert scaffold materials as well as (2) the knowledge on bioactive materials and is culminating in (3) the development of regenerative and custom-made biosynthetic implants and tissue grafts. It is the aim to fabricate scaffolds for engineered tissue units using biodegradable soft and simultaneously porous materials that allow the embedding and integration of biological cells with growth and differentiation factors, exogenously added or synthesized by the cells themselves.
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2.1 Bioinert Materials The application of bioinert materials for tissue and bone tissue replacement is old. Especially the application of biomaterials from marine animals as plaster and in bone/tissue replacement has been a tradition since the Greek times. Camper (1771) described that the organic matrix of sponges can be successfully used in plastic surgery of the palate of the skull (Figure 2.1(a)). He was fabricating a nose using lime wood, covered it with a sponge and fixed it in the roof of the mouth via a silk thread which had been waded by small sponge slices. It might also be noted that the application of the siliceous skeleton of sponges as suitable scaffold onto which human stem cells can be seeded has recently been reported (Green et al., 2003). Basically two basis materials have been used to fabricate implants for bone reconstitution/reconstruction; (1) metals and (2) ceramics. In the first phase of those replacement supports, the intention for the use of those implants had been to stabilize the body at the position of the damaged bone (Chang et al., 1996). For decades metal implants have been used in orthopedics for mechanical skeletal repair. Those supports had to meet the challenge to strengthen the implant–bone interface and to prevent stressshielding effects. Those implants can be fabricated in customized processes, e.g., by 3D printing (Figure 2.1(b)). The key issue for a durable and successful implant is the establishment of a strong bone–implant interface. It emerged that smooth implant surfaces can result in the formation of encapsulation with the consequence of loosening of the implant (Greco, 1994). In one approach to extend and to promote long-term interface strength, porous materials and porous coatings have been developed (Engh and Bobyn, 1985). Those porous materials and coatings induce a partial to complete bone ingrowth, which has the advantageous property to enhance the strength of the interface bonding under simultaneous reduced tendency to cause capsule formation around the implant. A further challenging issue for most of those metalbased implants is the appearance of “stress shielding” (Jacobs et al., 1993; Amstutz, 1991). Even though it is well established that bone regeneration and repair processes are promoted by mechanical loads (Van Lenthe et al., 1997) metal materials such as titanium, still widely used today for bone implants, is much stiffer than native bone. Consequently, an implant of solid titanium can carry a disproportionate amount of the biological loads (see: Thelen et al., 2004). In turn, the surrounding bone undergoes a process of “stress shielding” and suffers from abnormally low levels of
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A
B
C
D
Figure 2.1 The development of bone scaffolds. (a) Bioinert materials, e.g., application of sponges in tissue replacement (Camper 1771). The damaged nose/nostrils (A, B, D, C) was modeled by a piece of lime wood covered with a sponge (T, U, V), and fixed in the roof of the mouth (W) via a silk thread (S) that had been surrounded with small sponge slices. (b) Computer-aided rapid prototyping/3D printing. (A) Data are generated for an organ or tissue unit using the computing process. Algorithms for the automated design and fabrication of a scaffold/bone part are developed based on an assembly-free process. (B to D) The bone unit is printed, like in an ink-jet printer, and comprises a tightfitting customized morphology. If it is formed of ceramic or titanium, the implant has usually only an osteoconductive property. (c) Fabrication of an osteoinductive scaffold made of glass, metal, or ceramic that allows the cells to migrate into its pores. The scaffold is coated with bioactive factors or polymers, e.g., bone morphogenetic proteins-2 or polyphosphate, that direct its associated cells, stem cells, to terminally differentiated cells that build tissue units like blood vessels.
stress, which finally results again in bone resorption, followed by loosening of the implant (Black, 1999). A further achievement toward a more advanced material to be used for orthopedics is the use of bioinert ceramics (Boutin, 1981). Ceramic, an
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inorganic, nonmetallic solid that is prepared by the action of heat and subsequent cooling, has been used since the 1970s in orthopedics (Smith, 1963; Eyring and Campbell, 1969). As material, ceramics could have a crystalline or partly crystalline structure, or be amorphous (e.g., a glass). Ceramics are increasingly used in the orthopedic surgery, especially for joint prostheses. It is used not only for joint bearings, but also for the bone–implant interface of prostheses (see: Hayashi et al., 1993), especially for implants as knee prosthesis, total ankle prosthesis, or total elbow prosthesis. The bioinert ceramics have been found to have the property of excellent resistance during carrying (Kumar et al., 1991). Since the beginning of the development of ceramic implants it is hoped that this material is more biocompatible than metal alloys since it is provided with the property of resistance to corrosion, to be less cytotoxic and to be hydrophilic. At present synthetic biodegradable polymers, interconnected with porous calcium hydroxyapatite (HA) ceramics have been found to be very suitable composite materials for implants, since they can be combined with growth and development factors in a carrier/scaffold system, e.g., with recombinant human bone morphogenetic protein-2 (rhBMP-2), that strongly promotes the clinical effects of rhBMP-2 in bone tissue regeneration (Bianco and Gehron, 2000). This study, like others that appeared during this time period (Hollister, 2005), predetermines the development to a second level of regenerative restorative implantology the “bioactive materials.”
2.2 Bioactive Materials The strategy to design more effective bioactive tissue engineering scaffolds is to implement in a complex manner three essential elements, first, a porous matrix (scaffold), second, to elicit osteoconductive signals, and third, to implant, adjacent to the matrix, osteogenic cells that can attach to the matrix and respond to their signals via an adequate blood supply (Finkemeier, 2002); Figure 2.1(c). In turn, the implants must be designed in a hierarchical way; first hierarchical porous structures must be prepared that are provided with a suitable mechanical stability and flexibility and allow the transfer and diffusion of growth factors and differentiation factors. The complex 3D anatomical shape of the bone substitution material must try to imitate from the nanometer to the millimeter level the functional properties of the natural bone.The scaffold formed must ideally meet the requirements of the cells to be provided with nutrients. The porous channels must allow cell migration, and their surface features must be suitable for cell attachment (Cukierman et al., 2001). The morphology/topography as well as the
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roughness of the scaffold surface should fall between the theoretical maximum given by the material and the theoretical minimum of zero predicted by composite theories (Torquato, 2002; Hollister, 2005). Consequently, the critical issue for the design of the surface texture is based on computer calculations that match the requirements to allow intercellular and transcellular signal transmission as well as to leave space for the development of a vascular system that allows mass transport. An effective permeability is determined by the 3D pore arrangement and the adhesion receptors associated with the plasma membrane, e.g., the integrins. The biological properties of the surfaces of the, basically inert, matrices are crucially important.The material must be bioactive along an increasing complexity and inducibility, from being osteoconductive to osteoinductive and allowing processes that cause osseointegration. According to the definition by Albrektsson and Johansson (2001) the term osteoconduction means that bone grows on a surface, e.g., bone surface, that supports the ingrowth of the osteoblasts into pores, channels, or pipes (Wilson-Hench, 1987). However, more often the surface used is not the bone but another biogenic growth platform (Glantz, 1987). The next higher level of bioactivity is osteoinduction. This term stands for the activity of a contact or soluble material that displays the potency to induce the undifferentiated and pluripotent stem cells to enter the differentiation pathway toward the bone-forming cell lineage, more specific to induce osteogenesis. Finally, osseointegration refers to the process by which the implant is stably anchored into the bone. At present and strictly speaking, only the bone morphogenetic proteins (BMPs) being members of the transforming growth factor-β (TGF-β) superfamily of growth factors and well established physiological regulators of osteoblastic differentiation (Lavery et al., 2008) can be considered as osteoinductive. Among them BMP-7 and BMP-2 display the highest osteoinductivity (e.g., Urist, 1965; Lavery et al., 2009; Sampath et al., 1990). Already in limited clinical use is BMP-2 that has been shown to induce new bone formation in spine fusions and long bone nonunion fractures (Gautschi et al., 2007). After binding of BMP to the integrated cell surface receptor, a tetramer of serine/threonine kinase transmembrane receptors consisting of two type I and two type II receptors, intracellular signaling occurs via intracellular signaling proteins to the receptor-regulated Smads (R-Smads). In turn R-Smads form heteromeric complexes with the common mediator Smad, Smad-4, and subsequently translocate to the nucleus where they act as transcription factors to induce the BMP responsive genes (Sebald et al., 2004). This implies that any kind of scaffold supplemented/coated with BMPs must be qualified to be
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an osteoinductively acting implant. However, recently it becomes overt that in osteoblasts also BMP-independent anabolically acting routes exist. Examples are polyphosphate (polyP) and biosilica-mediated pathways (Wang et al., 2014a,b), natural inorganic polymers that have the capacity to induce BMP-2 in osteoblasts. In addition, a BMP-independent route has been proposed that induces the differentiation of osteoblast precursor cells to mature functionally active osteoblasts (Müller et al., 2013a). It also appears likely that the phytoestrogen isoquercitrin acts synergistically with polyP, on the transcription factor RUNX2 (Wang et al., 2014c). In line with a broader interpretation of osteoinductivity (Amini et al., 2012) biomaterials, including natural and synthetic ceramics (i.e., HA and various calcium phosphate (CaP) compositions, and their composites), have been qualified as osteoinductive materials. Besides of synthetic CaP-based biomaterials, also in the form of sintered ceramics (Yamasaki and Sakai, 1992; Klein et al., 1994), cements (Habibovic et al., 2008), and coatings (Habibovic et al., 2004), natural and coral-derived ceramics (Ripamonti, 1991; Ripamonti et al., 2009) have been attributed to be osteoinductive. Especially highlighted should be porous bioglass that has been widely introduced into clinics (Jones, 2013). This bioactive ceramics, a biodegradable glass of a general formula “Na2O–CaO–SiO2–P2O5,” contains high levels of calcium (Ca2+); its most generally used formulation contains 46.1 mol% SiO2, 24.4 mol% Na2O, 26.9 mol% CaO, and 2.6 mol% P2O5 and has been termed 45S5 Bioglass (Hench et al., 1971). The bioactive glasses are reported to stimulate bone regeneration to a larger extent than other bioactive ceramics. This 45S5 Bioglass forms a semichemical bond with bone and, in vivo, bonds to other bioceramics. In comparison, the CaP-based materials are more widely used in the clinics.
2.3 Regenerative Functional and Custom-Made Tissue Units Already a decade ago it has been prognosticated that tissue engineering technology, based on computer-aided jet-based 3D organ/tissuelike printing, could be a solution of the organ transplantation crisis (Mironov et al., 2003). However, it is needed to distinguish between inert implants that cannot be replaced by cells/tissue of the recipient and functional graft substitutes that are disintegrated and subsequently replaced by cells and ECM filaments ingrowth from the surrounding tissue of the recipient. Many types of CaP biomaterials have been developed that comprise a similar composition like native bone mineral and its precursors, e.g., HA and α- and β-tricalcium phosphate in the form of ceramics, cements, and thin
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coatings (LeGeros, 2002). A few of these insoluble CaP materials are even osteoconductive and in some cases provided with the ability to induce bone formation, implying that those materials are osteoinductive (Damien and Parsons, 1991; Habibovic and de Groot, 2007). The implant, even though acting to a limited extent osteoinductively, remains as a major core in the previously damaged region.The process of biodegradation around the bone graft substitutes is favored since they prevent the disadvantageous resorption of the neighboring bone due to stress-shielding effects. The technique of free-form fabrication applicable for producing of 3D synthetic bone graft substitutes allows a precise control of the overall geometry and in turn also of the porous structure of the scaffold/implant. The 3D printing of CaP-based structures at room temperature has been successfully demonstrated (Gbureck et al., 2007). By modification of the surface of the implant, either with respect to the morphological, ultrastructural, or chemical properties, the capacity of the material to induce ectopic bone formation can be improved. The fabrication of 3D implants by direct cell printing based on computer-aided design files offers a sophisticated and challenging direction to engineer 3D tissuelike units, to be placed into living human organs (Figure 2.2). However, those approaches require three sequential hierarchical steps of increasing complexity. First, compilation of preprocessing followed by the development of computational “blueprints” of a given organ; second, processing or the program file for the printing of the actual organlike unit; and finally, postprocessing or organ conditioning and ultimately organ maturation. In recent years several types of cell printers have been developed (Tasoglu and Demirci, 2013). The basic principle is that the cells, either in suspension or as aggregates, are embedded into a printable matrix which is then sequentially layered under formation of predesigned blocks. It is surely feasible to fabricate implantable, by rapid-prototyping, 3D organlike units in the future; but it remains open when this concept of tissue engineering can be exploited and integrated into the constraints emerging from the biological, genetic rules of developmental biology. It is just extraordinarily difficult to tailor a suitable matrix into which the cells can be embedded to become provided with the required physiological solute and fibrous extracellular molecules. Only after deciphering the genetic blueprint of the cells and its time- and space-specific expression it will become possible to lay the ground with the chemical and physical cues for the stem/precursor cells to direct them toward an integrative assembly in an organ.
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Figure 2.2 Regenerative functional and customized tissue units bioprinted by the 3D cell printing approach. Sketch of the bioprinting of cells, embedded into a matrix, e.g., alginate/gelatin, using a 3D bioplotter. A cell suspension is filled into a cartridge hooked to the printing head (a). This control element is connected with the computer-guided printing apparatus; the alginate/gelatin/cells are passed through a needle into a CaCl2 bath, which hardens the scaffold (b, c). This scaffold with the bioprinted cells is submersed into medium/serum. Then the 3D scaffold is overlayed with an agarose layer containing the morphogenetically active factors or polymers, e.g., polyP (d). In such an environment the cells proliferate and differentiate.
For 3D bioprinting a critical size of the aggregates has to be intended. A cell density within the organic matrix of >106 cells/ml, according to our experiences, is preferable (Neufurth et al., 2014). The size of the aggregates formed within the scaffold matrix is determined by the supply of nutrients and growth factors and/or morphogenetically active polymers. Usually the aggregates are spherical and can reach sizes of a diameter of 500 μm. Incubation conditions must be developed that are favorable for the cells to spread onto a cell substrate, e.g., fibronectin or collagen. In turn, cell–cell and cell-substrate adhesion must be finely tuned; the cells must express the property to decrease substratum adhesivity while simultaneously allow increasing cell–cell cohesivity, and vice versa (Ryan et al., 2001). In addition, the cell environment should be tailored in a way that one cell type is directed towards cell–cell cohesion while the other cell type undergoes intensive cell spreading. The cell aggregates formed after bioprinting can be composed homocellularly, comprising only one type of cell or heterocellularly, being a hybrid composed of more than one cell type (Figure 2.3(a) and (b)). If homocellular aggregates are formed in a bioprinted matrix the subsequently formed
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Figure 2.3 Bioprinting of cells, embedded in an organic matrix. (a) Either homocellular aggregates are allowed to be formed that have the potency to form basic but simple building units. Alternatively, (b), heterocellular aggregates are allowed to be formed that can differentiate to functional units, e.g., comprising vascular structures (v), within cells derived from a different stem cell lineage. (c) It is still an ambition to bioprint cells of different stem cell origin that can functionally differentiate to different cell types that build an organ.
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tissue structures can show different morphologies but hardly form functional organ units (Figure 2.3(a)). In contrast, if heterocellular aggregates are allowed to be formed within the matrix the formation of internal structures can be expected, e.g., blood vessel within endothelial and smooth muscle cells. Alternatively, peripheral blood stem cells can be used that can differentiate either into endothelial cells or smooth muscle cells; endothelial progenitor cells are present in peripheral blood stem that can be induced to cells supporting neoendothelialization (Sugaya et al., 2012). At present it seems to be manageable to sequentially bioprint cell types with a different phenotype (Figure 2.3(b)). The solution of choice in the future will be to bioprint cells that, based on their cell adhesion properties and origin from a given stem cell lineage, proliferate and in parallel differentiate to a complex organ (Figure 2.3(c)).
3. BONE CELLS Bone appears to be a solid, rigid organ; however, it is highly flexible and dynamic allowing bone anabolic and bone catabolic processes to proceed in a tuned interacting manner. The bone is under continuous remodeling that takes place throughout the life span of an individual. It should be highlighted that the morphology and the shape or the size of bone is genetically determined. Until now it remains unclear which genetic blueprint controls the form-giving processes in bone. Under physiological conditions the net balance between osteoblastic bone formation, mediated by osteoblasts, and osteoclastic bone resorption, driven by osteoclasts, is very much tuned. The bone anabolic cells, the osteoblasts, originate from MSCs having the potential to proliferate and the capacity to differentiate into several connective tissue/cell types. In contrast, the bone catabolic cells, the osteoclasts arise from hematopoietic stem cells (Teitelbaum, 2006).
3.1 MSCs/Osteoblasts The pluripotent MSCs have the potency to differentiate into osteoblasts, chondroblasts, bone marrow stromal cells, fibroblasts, muscle cells, or adipocytes depending on the presence of the growth and differentiation factors in their microenvironment (Wang et al., 2014a); Figure 2.4. Osteoblasts having a cuboidal or columnar shape are lining the bone surfaces at those sites that undergo active bone formation during bone development or fracture repair. Osteoblasts express high levels of type I collagen (COLI) and proteoglycans (glycosaminoglycans), the two main components of the bone matrix, also
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Figure 2.4 Multipotent differentiation of multipotent human mesenchymal stem cells (hMSC). Specific signaling molecules and growth factors as well as differentiation factors induce/activate transcription factors and by that determine both the commitment and the differentiation of hMSCs toward the osteogenic, chondrogenic, adipogenic, or myogenic lineage. The osteogenic and the chondrogenic lineages are involved in the restorative repair of bone and cartilage tissue (osteochondral tissue reconstitution). Biosilica and polyphosphate (polyP) display anabolic, morphogenetic effects on those two differentiation lines. BMP-2, bone morphogenetic proteins-2; ALP, alkaline phosphatase.
termed osteoid. Osteoblasts are also involved in mineralization of osteoid, very likely via the liberation of matrix vesicles, and by the deposition of calcium, carbonate, and phosphate (Landis et al., 1993; Hohling et al., 1978; Müller et al., 2013b). Osteoblasts are aligned by adherens-type junctions, including desmosomes and tight junctions (Safadi et al., 2009). Osteoblasts synthesize and secrete a variety of cytokines and colony-stimulating factors controlling myelopoiesis, e.g., interleukin-6, interleukin-11, granulocyte–macrophage colony-stimulating factor and macrophage–colony-stimulating factor
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(M-CSF). In addition, osteoblasts synthesize a series of growth factors, including TGF-β, BMPs, platelet-derived growth factors, and insulin-like growth factors. Finally, terminally differentiated osteoblasts possess receptors for the parathyroid hormone as well as for 1,25-dihydroxy-vitamin D (1,25(OH)2D3), the major hormones regulating bone metabolism and mineral deposition (Figure 2.5).
3.2 Osteoclast Cell Differentiation Osteoclasts, originating from the hematopoietic lineage (Boyle et al., 2003), undergo differentiation and maturation in the presence of the M-CSF and of the receptor activator of NF-κB ligand (RANKL). Markers for the multinucleated osteoclasts are the highly expressed tartrate-resistant acid phosphatase (TRAP), as well as calcitonin receptor and integrin avb3 (Cerri et al., 2003); Figure 2.5. The cytokine/receptor triad, RANKL with its receptor (RANK) and the endogenous decoy receptor osteoprotegerin (OPG) crucially control bone formation and bone remodeling (Boyce and Xing, 2008; Santini et al., 2011). While RANKL is synthesized by the osteoblastic lineage cells this signaling molecule is essential for the differentiation of those cells that are involved in bone resorption, the osteoclasts. RANKL is expressed on osteoblasts, T cells, dendritic cells, as well as their precursors from where it can be released by specific proteases (Zhang et al., 2009).This ligand (RANKL) binds to the cell surface receptor RANK, located on precursor and mature osteoclasts and by that promotes osteoclastogenesis. After binding of RANKL to RANK the osteoclasts become activated and resorb bone mineral; during this process the cells have close contact to the bone surface (Fuller et al., 2010). At this interphase, osteoclasts to bone, vesicles are formed, via integrin (avb3), that contain proton pumps and acid hydrolases, e.g., cathepsin K. Those enzymes and vesicles are present in those cells that are bone-apposed. “Resorptive hemivacuoles” are formed between osteoclasts and bone, allowing the protons to dissolve the HA scaffold of the bone (Figure 2.5). The intracellular pH is kept close to neutral via the chloride/bicarbonate exchanger.The function of RANKL is under control of OPG, a decoy receptor that is secreted by stromal cells and also by osteoblasts (Kearns et al., 2008). OPG scavenges RANKL by binding to it and neutralizes its function. In turn it has to be concluded that any deregulation of the tuned expression of the RANKL/RANK/OPG system causes a dysregulation of the differentiation pathways of the osteoblasts and the osteoclasts and in turn promotes catabolic bone remodeling (Boyce and Xing, 2008). By that, OPG prevents bone matrix from excessive
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Figure 2.5 Differentiation of the progenitor cells of the bone-forming osteoblasts (osteoblastogenesis) and the bone-resorbing osteoclasts (osteoclastogenesis). Upper panel: Osteoblast differentiation starts from the mesenchymal stem cells and ends with the osteocytes. The major transcription factor Runx2, which is under the control of bone morphogenetic proteins-2, is synthesized in chondrocytes and causes a stage- dependent increase in the structural and functional proteins in osteoblasts, for example, b-ALP (bone-specific alkaline phosphatase), COLI (collagen type I), OP (osteopontin), ASP (asialoprotein), BSP (bone sialoprotein), and OCAL (osteocalcin), as well as RANKL (receptor activator of NF-κB ligand). Lower panel: Principle differentiation stages from the hematopoietic stem cells via preosteoclasts to functionally active, bone-resorbing osteoclasts. The osteoblasts direct the preosteoclasts to the osteoclast through the interaction of RANKL with RANK (receptor activator of nuclear factor κB), an interaction that is blocked by OPG (osteoprotegerin). Differentiation from hematopoietic stem cells starts via activation of the PU.1 transcription factor and inflammatory signals. The CD34+ osteoclast precursor cells, after entering the circulating system and in the presence of M-CSF (macrophage–colony-stimulating factor) and 1,25-dihydroxy-vitamin D3 (vitamin D3), become recruited onto the surface of bone. The preosteoclasts, after the stimulation of the DAP12 adapter protein/receptor undergo multinucleation to the osteoclasts. Those cells express in the presence of 1,25-dihydroxy-vitamin D3 the receptor RANK. After binding of RANKL to RANK the osteoclasts dissolve HA by lowering the pH. Markers for the activated osteoclasts are TRAP (tartrate-resistant acid phosphatase) and CTR (calcitonin receptor). HA, hydroxyapatite.
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resorption by binding to RANKL and in turn abolishes the activation of the osteoclasts via inhibition of the RANK pathway. In conclusion, the relative concentrations of RANKL and OPG in bone are the major morphogenetic determinants of bone mass and strength.
4. BIOGENIC, MORPHOGENETICALLY ACTIVE INORGANIC POLYMERS 4.1 Biocalcite Exemplarily the beneficial function of the calcareous corals in bone reconstruction has been demonstrated (Cooper et al., 2014). Especially the effects of secondary metabolites from soft corals acting against inflammation and tumor growth have been highlighted (Chen et al., 2013). Well understood is the inductive osteogenic differentiation effect of coral scaffold on MSCs (Puvaneswary et al., 2013). Even, with respect to some biological markers, calcareous scaffolds, derived from corals, have been found to be superior in comparison to bone grafts. A significantly higher level of expression of the osteogenic differentiation markers, ALP, osteocalcin (OCAL), and osteonectin, as well as of the transcription factor Runx2 has been described. Even more, the extent of mineralization within coral grafts has been found to be more extensive compared to bone grafts. Further studies revealed that coral products have a curative potential on bone deficits as well. The naturally occurring calcium, within the calcareous scaffold, in the form of aragonite found in the scleractinian hard corals and in the form of calcite deposits within the soft octocorals, contribute to anabolic bone restoration. This effect is especially pronounced if these minerals are administered together with zeolite, a microporous mineral, in mice induced to a menopausal state (Banu et al., 2012); this effect has been confirmed in rabbits as well (Parizi et al., 2012). Parallel with the effect of coral minerals on bone formation, their effect on dental deformities has been studied (Figueiredo et al., 2010). The data indicated that the coral skeleton in its unrefined form cannot be applied due to its necrotic potential. However, if purified coral minerals are applied a beneficial osteogenic effect on bone marrow stromal cells is seen; an effective repair of mandibular defects in canines has been reported (Yuan et al., 2010).
4.2 Bio-polyphosphate Like the chemically synthesized inorganic polymeric phosphate, polyP, the biogenic polyphosphate (bio-polyP) has an amorphous state (Kulaev et al.,
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2004; Rao et al., 2009; Omelon and Grynpas, 2008). In contrast to the chemically synthesized polyP, which is synthesized at high temperature, the biogenically produced bio-polyP is synthesized at ambient physiological conditions (Rao et al., 2009) via polyP kinases; Figure 2.6(a). The biopolymer bio-polyP is found in a wide range of organisms, including bacteria, fungi, algae, plants, and animals (see: Rao et al., 2009); it is readily water soluble in millimolar concentrations at chain lengths
