Calcium Phosphate Scaffolds: History
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Contributor:
Federico Pupilli
,
Andrea Ruffini
,
Massimiliano Dapporto
,
Marta Tavoni
,
Anna Tampieri
,
Simone Sprio

Bone is a complex biologic tissue, which is extremely relevant for various physiological functions, in addition to movement, organ protection, and weight bearing. The repair of critical size bone defects is a still unmet clinical need, and over the past, material scientists have been expending efforts to find effective technological solutions, based on the use of scaffolds. In this context, biomimetics which is intended as the ability of a scaffold to reproduce compositional and structural features of the host tissues, is increasingly considered as a guide for this purpose. However, the achievement of implants that mimic the very complex bone composition, multi-scale structure, and mechanics is still an open challenge. Indeed, despite the fact that calcium phosphates are widely recognized as elective biomaterials to fabricate regenerative bone scaffolds, their processing into 3D devices with suitable cell-instructing features is still prevented by insurmountable drawbacks. With respect to biomaterials science, new approaches maybe conceived to gain ground and promise for a substantial leap forward in this field. 

  • biomimetics
  • hydroxyapatite
  • 3D scaffolds
  • bone regeneration

1. Introduction

The skeleton plays a fundamental role in human physiology for both protection and support of many vital organs. At the same time, it serves as a framework for the body, providing attachment for the muscles and fibrous connective tissues, such as ligaments and tendons. Bone tissue may be subjected to trauma or other degenerative diseases during lifetime. Indeed, the regeneration of impaired bone tissue is still a largely unmet problem, particularly when it comes to the treatment of critical size and load-bearing bone defects, which implies the occurrence of serious disabilities and impact on the human wellbeing as well as on direct and indirect healthcare costs [1]. Clinical treatment of bone defects typically involves bone grafting techniques in orthopedic, cranio-maxillofacial, and spinal surgery. For years, the use of bone autografts was considered as the most suitable choice for the purpose of bone repair and regeneration, in terms of histocompatibility and non-immunogenic properties. Indeed, the harvest of bone from the patient consists primarily of cortical and cancellous bone with the essential components to achieve bone regeneration processes, such as growth factors including bone morphogenetic proteins (BMPs) and osteoprogenitor cells. Conversely, the use of allografts is typically associated with risks of immunoreactions and transmission of infections. Due to the devitalization processes (irradiation or freeze-drying processes) following the extraction from the deceased donor, allografts have reduced osteo-inductive properties and no cellular component [2]. On the other hand, however, autografts require generally a second intervention at the site of the harvest, resulting in a very expensive and painful surgical practice associated with surgical risks, such as infection and chronic pain [3][4]. For these reasons, decades of intensive effort are dedicated to the development of synthetic bone grafts with effective regenerative ability while mitigating the risks associated with clinical treatment [5].

2. Translation of the Biomimetic Concept to 3D Scaffold Development

2.1. Limitations of Current Approaches and Further Challenges in Tissue Engineering

As already stated, one of the main targets in the design of biomimetic CaP ceramic scaffolds is the ability to create a micro-environment that stimulates cell differentiation into osteoblasts and to stimulate cell chemotaxis and new bone matrix deposition. Although HA has many advantages in terms of bioactivity and bioavailability, it is limited by fragility, typical of ceramic materials, poor mechanical strength, and thus, inability to withstand severe load-bearing conditions. One of the most common ways to confer mechanical properties to CaPs (and ceramic phases in general) is via densification processes obtained through high temperature sintering processes. Although densification usually yields ceramic bodies with superior mechanical strength, the properties of CaP materials in terms of crystallinity, grain size, porosity, and composition vary significantly upon sintering. High crystallinity, low porosity, and small grain size tend to provide higher stiffness, compressive strength, and toughness, but hamper tissue growth and binding of bone-resorbing cells [6].
Sintering temperature of synthetic HA in a range between 950–1150 °C causes a general increase in grain size, density, crystallite size. Upon heating to 1250 °C, samples undergo several structural transformations with the formation of α- and β-tricalcium phosphate (TCP: Ca3(PO4)2) that, although biocompatible, are not found in bone [7]. The situation further complicates when non-stoichiometric HA, but rather calcium-deficient or biomimetic ion-doped HA, are considered. Studies conducted on Mg- and Sr-substituted HA prove that the introduction of Mg ions, at a weight percentage of 0.5% provokes the formation of both β-TCP and CaO upon thermal treatments, in contrast with what occurred in pure HA, which did not dissociate in secondary phases. These phenomena could be asserted to the distortion effect of the crystal lattice caused by the substitution of calcium with foreign atoms, which ultimately may lead to dissociation in secondary phases upon thermal treatment, with the amount of secondary phases increasing by raising the working temperature [8][9].
Problems related to the formation of secondary phases during the densification processes, in association with reduced biological activity, substantially limit the regenerative effectiveness of sintered scaffolds in bone tissue engineering. In the last decades, studies have been conducted to explore sintering behavior under microwave heat treatments rather than conventional sintering. Microwave-based approaches allow for the attainment of consolidated scaffolds with controlled structure, high densification extent, and fine grains, with a significant increase in mechanical strength with respect to conventional sintering, in addition to a higher solubility with respect to conventional sintering, leading to significant cell adhesion, distribution, and proliferation in vitro [10].
Another common strategy that has been largely used in bone scaffold development is the achievement of porous scaffolds with a 3D arrangement blending CaPs in addition to biodegradable polymers. In this context, 3D printing has become a widely used technique to fabricate composite scaffolds in regenerative medicine and is defined as a computer-aided transfer process to pattern and assemble materials with a prescribed 2D or 3D organization to fabricate bio-engineered structures with precisely designed macro-architectures [11]. Apatite powders on their own do not possess consolidation behavior that enables printability, thus it is necessary to optimize bio-inks to achieve final devices with appropriate mechanical integrity through the combination of CaP powders with polymer materials, as they possess sufficient mechanical properties and are suitable for the repair of critical bone defects [12]. In this latter aspect, various biocompatible polymers, such as polylactic acid (PLA), poly(lactic-coglycolic acid) (PLGA), and polycaprolactone (PCL) have been used for the fabrication of bone implants and even received approval from the US FDA as materials for 3D printing of biomedical implants [13]. Among them, polylactic acid (PLA) has been defined as a biomaterial with potential clinical applications in many studies due to its slow degradation properties and reliable biocompatibility [14]. Studies have been focused on the fabrication of composite nano-HA/PLA and characterization of the mechanical properties, in vitro biocompatibility, and in vivo experiments in a rabbit femoral defect model for 3 months, showing good biocompatibility and osteogenic induction ability by simulating organic and inorganic components of bone tissue, and with the potential to repair critical bone defects [15]. Nevertheless, although promising in terms of biocompatibility, one main concern of the use of these polymers in bone tissue engineering is that their intermediate degradation products (specifically, lactic acid and/or glycolic acid) by non-enzymatic hydrolysis of ester bonds in their backbone reduces the local pH, which in turn can induce inflammatory reactions and bone cell damaging at the implant site [16]. Moreover, the rapid drop of pH in vivo may accelerate the polymer’s degradation rate, thus resulting in the dispersion of inorganic particles in vivo and also in the premature loss of mechanical properties before new bone formation occurs. Another drawback related to the use of polymers added with inorganic particles is the possible loss of plastic behavior under loading, in order to be ineffective in terms of reinforcement and actually result in a decrease in fracture strength [15][17][18].

2.2. Guiding Bone Regeneration by Chemistry and Crystal Structure

Hydroxyapatite is the most stable and the least soluble CaP and it is considered as the main inorganic model for the design of suitable biomedical materials due to its similarities to bone and tooth enamel. As previously stated, the crystal lattice of apatites is characterized by a loose-packed structure that permits the accommodation of foreign atoms in replacement of Ca and P. This generates surface irregularities and structural defects that ultimately result in an increased biosolubility due to the fact that the hydrolysis rate increases at lower Ca/P ratios [19][20][21]. The accumulation of defects generates lattice distortions and describes why the crystallinity of ion-doped HA is usually, yet not always, poorer than its pure analogue. The lattice disorder induced by ionic substitutions is the key factor at the basis of enhanced functionalities of biomimetic apatite phases, enabling the increase in the surface charge and consequent ability of prolonged exchange of bioactive ions, such as Ca2+, Mg2+, and HPO4 2− with the surrounding biological environment, which triggers bio-specific chemical signals active in bone regeneration [22].
For the above reasons, a main strategy to activate bioactive chemical signaling that mimics the ability of bone mineral is the achievement of low-crystalline, bioactive apatites by the introduction of doping ions into its crystal lattice. In the design of biomimetic scaffolds, Ca2+ sites may be occupied by divalent or monovalent cations, such as Na+, K+, Mg2+, Sr2+, Zn2+, Ba2+, Cu2+, and Fe2+, whereas PO43− could be substituted by atoms, such as CO32− and SiO44−, while hydroxyl (OH−) may be replaced by CO32−, F−, Cl− or even be left vacant [23].
Among all substitutions, the most prominent and studied in both biological and biomimetic tissues are the effects of carbonation. Biological tissues, such as bone, dentin, and enamel express a substantial difference in carbonate content, influencing their solubility in a biological environment. Mineral structures present in dental tissue, for example, introduce a crystal strain and increase solubility, thus dentine mineral is considerably more soluble than enamel mineral and both are considerably more soluble than stoichiometric HA [24]. The considerable differences in structure between dentine and enamel influence the interactions of the tissues with acid solutions, thus the relative rates of dissolution do not necessarily reflect the respective solubilities [25].
CO32− ion substitution into the apatitic crystal lattice has been shown to considerably enhance the dissolution rate and improve the solubility behavior compared with stoichiometric HA [24][26][27]. In vitro studies have found an increased collagen production by human osteoblast cells on carbonated HA compared with undoped HA, as the secretion of type I collagen strongly depends on extracellular calcium concentrations that resulted in a remarkably higher cell culture medium containing carbonated HA [28].
Magnesium plays a key role in bone mineralization and bone as it improves the osteointegration and osteoblasts activity and accelerates bone ingrowth, enhancing integrin ligand binding and protecting cells by oxidative stress [29][30]. The presence of Mg2+ ions on the apatite surface results in an increase in the amount of water molecules adsorbed on the surface, affecting cell attachment and differentiation [31][32].
The incorporation of Sr2+ ions into the apatite lattice results in an increase in the osteointegration process. It was demonstrated that the incorporation of Sr2+ ions into the apatitic crystal structure promotes osteoblasts differentiation and proliferation, enhancing the activity of alkaline phosphatase (ALP), as well as the production of collagen type I and osteocalcin [33]. Thanks to the ability to reduce the proliferation of osteoclasts by hindering their activity, the presence of Sr2+ in the apatite lattice is directly linked to the treatment of osteoporotic patients in case of tumor surgery incision or trauma [34][35]. In addition, studies have shown that the partial substitution of Ca2+ with Sr2+ can stabilize the apatitic crystal lattice, also enhancing the scaffold hardness and the mechanical stability of the newly formed bone [36][37]. Although in tissue engineering ion-doped apatites with tailored crystal structure can be very relevant in mimicking the behavior of biological apatite, ion-doped apatites in the form of nanopowders, which can be usually obtained in laboratory, are poorly useful in bone regeneration procedures, particularly when large bone defects are to be treated. In this context, a further step is required in the development of 3D scaffolds to retain cell-instructive ability and appropriate porous structure and mechanical properties.

2.2.1. Synthesis Processes for the Production of Biomimetic Apatites

The use of HA in clinical application has been extensively investigated for more than four decades. HA is commercially available from a natural source or as synthetic product. Various methods were developed for the synthesis of crystalline HA phases, such as solid–state reactions involving the mixture of β-TCP and Ca(OH)2 powders in specific ratios in water, followed by wet milling, casting into bodies, drying, and sintering at high temperature (typically at least 1000 °C). However, the use of high temperature during these sintering treatments makes this method not ideal for the achievement of biomimetic, nanocrystalline, and bioactive apatites [7][8][9]. Therefore, various wet processes have been developed in the last decades, such as precipitation, hydrothermal, and sol-gel methods. With these methods, physical and morphological characteristics of synthetic apatites can be tailored by regulating the process conditions during the synthesis (pH, reaction time, temperature, concentration, type and state of precursor, etc.) [34][38].

2.2.2. Tailoring the Dissolution Mechanism and Solubility of Apatites

As previously mentioned, bone modeling–remodeling processes take place after the attachment of osteoclasts on the surface of bone tissue/implants, followed by resorption occurring through dissolution of the inorganic phase and the enzymatic degradation of the organic bone matrix. When designing devices for bone regeneration, the ability to undergo osteoclastic resorption is a key property. The bone mineral is a major actor in this context, and thus it represents a model for material scientists, as chemical–physical parameters observed in biological apatites, such as crystallinity, foreign ion content, and crystal morphology all entail more effective and different solubility behavior, which in turn affect the response of both the osteoclasts as bone-resorbing cells and osteoblasts as bone-forming cells [34]. Multiple models have been formulated to exemplify biodegradation in bone remodeling processes, simulated by the dissolution of apatites in an acidic environment [39][40].
This simplistic model takes into consideration perfect crystalline structures with a smooth surface (dissolution steps are absent), where the detachment of one or several ions results in the formation of dissolution nuclei. Surface irregularities and structural defects all may act as a nucleus of the subsequent dissociation process [34]. As a way to mimic the resorption behavior of bone mineral, the study of biomimetic apatites with tailored properties has been widely studied through the insertion of structural defects upon ion substitution. In this context, some studies have investigated the solubility behavior of ion-doped and multi-doped apatites, considering that each ion has limited solubility in the apatite lattice, thus resulting only in partial substitution. Although the insertion of foreign bioactive ions, such as Mg2+, Sr2+, and CO32− ions in the apatitic structure causes a general enhancement of the overall solubility in synthetic biological fluids or in cell culture media [22][29][41][42][43][44][45][46][47], the solubility behavior of doped and multi-doped apatites can be quite difficult to predict. In fact, the extent of ion substitution in apatites is subjected to various and interlacing key aspects and the co-existence of different foreign anions and cations can influence substantially the solubility behavior of the resulting matrix. First, synthesis processes conducted at a relatively low temperature (such as body temperature) yield apatitic phases with limited crystal growth, which favors ion doping. On the other hand, ion substitution is facilitated when no charge imbalance occurs, such as, for instance, with divalent ions, such as Mg2+, Zn2+ or Sr2+ replacing Ca2+. Conversely, trivalent ions, such as Fe3+ or Ga3+ can be more easily hosted within the surface, non-apatitic layer [48][49]. Ions populating this disordered non-apatitic layer are likely less tightly bound to the apatite structure and can be more easily released and even exchanged within a fluid environment [16]. For instance, previous studies found that the simultaneous presence of Si4+ (replacing PO43−) and Mg2+ (replacing Ca2+) in synthetic apatites mutually limits their release in simulated body fluid (SBF) [50]. This phenomenon suggested the formation of a complex containing these two ions, but it was not possible to precisely determine its nature nor its location within the overall apatite structure. In a general way, the few studies conducted to date on the ion release from apatites converge on some aspects, such as the relevance to have B-site carbonation and to achieve multiple ion doping. However, more systematic approaches are required to elucidate the fine physicochemical mechanisms determining the ion doping extent and particularly to the mutual interaction between different ions competing to enter in the same crystal sites.

2.3. Guiding Bone Regeneration by 3D Scaffold Architecture and Porosity

Guiding bone regeneration first at the surface, and then in the bulk of synthetic scaffolds is one of the main challenges during current clinical treatments. In this context, it is necessary to take the 3D bone structure as an example and recreate its morphological features in synthetic scaffolds, controlling the surface structure and scaffold porosity. Bone macrostructure presents a substantial difference in porosity, progressing from the compact cortical bone, where porosity ranges from 5% to 10%, toward the inner cancellous bone, which is more porous, from 50% to 90% [51]. Studies conducted on HA scaffolds with parallel cylindrical pores of various sizes without interconnecting fenestration between adjacent pores showed good results in terms of osteoconduction at average pore size around ϕ50 μm, with optimum conditions at ϕ300 μm [52]. The overall porosity has a direct impact not only on cell attachment and their differentiation into bone reforming cells, but also on the rate of vascularization after implantation. Lack of pore interconnection in a 3D scaffold, for example, could inherently affect the overall bioresorbability of the scaffold itself, as bone-resorbing cells cannot penetrate and attach to the whole inner scaffold area and consequently remain only on the outer surface [53][54].
As fairly connected with osteoinduction and osteoconduction, the adherence of the implanted scaffold on the surrounding bone tissue is very relevant in tissue engineering. In this context, osteointegration refers to a phenomenon where an implant becomes intimately connected with bone in order to be unified as a whole, making this phenomenon necessary for long-term stability [55][56][57]. Porous apatitic scaffolds exhibit good mimicry of the bone mineral composition, promoting formation of new bone and a tight bone–implant interface within weeks from surgery. Highly porous structures further facilitate extensive bone penetration throughout the whole scaffold, with excellent osteointegration and positive effects on the overall biomechanical performance [58]. In vivo studies on mandibular sheep defects filled with HA scaffolds with defined macro-porosity proved that homogeneous, interconnected pores favored the formation of interpenetrating matrices of newly formed bone, thus leading to better integration and functionality of the construct [54]. The attainment of synthetic grafts with ordered, channel-like porosity from a clinical perspective may increase cell seeding efficiency and the distribution of viable cells in the inner part of the scaffolds by improving fluid conductivity and permeability, assisting in the prevention of necrotic regions formation [59]. Studies have proven that channel-like ordered porosity was found to promote the regrowth of osteon structures, whereas randomly-oriented porosity was more likely to favor the formation of new woven bone [52][60]. In addition, as the newly formed bone tissue needs nutrients and oxygen in order to subsist resorption processes, ordered interconnection is needed for internal vascularization. For this purpose, there are also other factors to take into consideration other than the overall porosity of the biomaterial, such as pore shape, interconnection, and arrangement, since they are possibly the limiting factor in bone osteoconductive processes and are determinants for optimal cell migration and blood vessel ingrowth [61].
It should be noted that, although higher porosity drastically increases bone reforming processes, biomaterials for implantations with high porous structures suffer from poor mechanical properties in terms of compressive strength and elastic modulus, which cannot meet the requirements of long-term orthopaedic applications [62][63]. For this purpose, material scientists are called to the difficult task of balancing osteoconductivity, vascularization ability, and mechanical performance by tailoring the porosity extent as well as the pore size distribution and interconnection to achieve mechanical strength, which is sufficient for early in vivo loading upon implantation and elastic properties close to those of bone. Then, osteogenic and osteoconductive ability should permit the formation of well-integrated new bone, thus progressively recovering the natural mechanical functionality of bone tissue. This is relevant to actively respond to complex biomechanical loads and activate mechano-transduction phenomena, a fundamental aspect guiding the formation and remodeling of new mechanically-functional bone [60]. All of these aspects will be considered in the next chapter, illustrating some recent approaches to develop biomimetic 3D bone scaffolds.

3. Recent Approaches Yielding Biomimetic Ceramic-Based Scaffolds

3.1. Organic/Inorganic Scaffolds by 3D Printing

Biomimetic strategies focus on the structure and components of bone tissue as models and recreate its features in synthetic scaffolds. The incorporation of natural polymers, such as collagen results in an ideal strategy in the development of bioactive organic/inorganic composite scaffolds. Three-dimensional printing techniques have been extensively investigated for this purpose. In previous studies, collagen was used in blends with biodegradable PLA and recombinant human bone morphogenetic protein-2 (BMP-2), showing comparable results with autografts for spinal fusion surgeries [64][65][66]. In another study, PCL embedding HA nanoparticles and blended with collagen showed good printability and cell proliferation proportional to the PCL content [67]. In addition, 3D printing was used to obtain biocomposites consisting of fibrillated collagen, tricalcium phosphate particles, and human umbilical cord serum (hUCS), with promising results in terms of osteogenic activity in vivo [68].
As previously mentioned, one of the main problems associated with the use of biodegradable polymers in the development of biomimetic scaffolds, such as PLA, PLGA, and PCL is related to degradation problems that could affect the overall osteointegration process. Furthermore, the highly viscous nature of polymers can lead to technological issues related to inhomogeneous infiltration, inappropriate pore interconnectivity, as well as a significant reduction in the overall porosity [69], which may hinder vascularization of the resulting composite material and also affect the final mechanical performance [70][71]. In this context, elastomer materials, such as poly(glycerol sebacate) (PGS) were recently investigated to more closely approach the mechanical properties of biological tissues [72][73][74]. Further introduction of polyethylene glycol (PEG) has been addressed to increase hydrophilicity and the resulting cell adhesion, proliferation, and differentiation on the scaffold surface. PEGylated PGS (PEGS) modified polymers have been infiltrated in CaP multiscale porous scaffolds leading to CaP/PEGS hybrid scaffolds, which still presented an hierarchically porous structure. Improved mechanical behavior and enhanced effects on in vitro cellular responses, as well as further in vivo experiments have been observed by simultaneously adjusting the polymer-coating amount and PEG incorporation in coating PGS [71]. In a further experiment, the incorporation of urethane-based PEGylated poly(glycerol sebacate) (PEGSU) in ceramic bio-inks results in freestanding hyperelastic bioscaffold, which is able to fix specific bone defects in craniomaxillofacial districts with good aesthetic results. Despite the fact that these polymeric components do not reproduce the biological features of natural polymers, such as collagen, their use can help in modulating the rheologic properties of bio-inks and facilitate micro-extrusion processes, yielding scaffolds with complex shape and geometry [75].

3.2. 3D Hybrid Scaffolds Using Natural Polymers and Bio-Inspired Mineralization Processes

Three-dimensional printing approaches are increasingly attempted, also pinning on the technological advances in 3D printing equipment. However, their use is still quite limited with natural polymers, particularly when containing mineral phases, due to the difficulty in obtaining appropriate rheological behavior. The use of various natural polymers in the development of composite scaffolds with enhanced bioactivity has been largely studied (for example, collagen, glycosaminoglycans, cellulose, and gelatin) [20].
Cellulose-based scaffolds were studied for their ability to nucleate bioactive calcium phosphate crystals in vitro, showing that cellulose scaffolds could be used as a model for in vitro studies [76]. In a different study, the electrospinning of gelatin blended with poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV) resulted in a biomimetic composite, which is able to form CaP crystals within the whole scaffold, with PHBV content strongly affecting its chemical composition and surface characteristics, in turn affecting the cytocompatibility [77]. On the other hand, gelatin modified with CaP nanoparticles and PCL was used to prepare a 3D bi-layer scaffold by collecting electrospun PCL and gelatin/CaP fibers separately in the same collector to enhance mineralization, thus improving the ability of the scaffold to bond to the bone tissue [78]. The use of PCL and CaP nanoparticles resulted in a synergistic effect: On the one hand, the PCL improved the mechanical properties of the scaffold, whereas the addition of CaP nanoparticles enhanced its bioactivity, which is attested by higher ALP activity with human osteoblast cells.
Previous studies highlighted the possibility of inducing the heterogeneous nucleation of bioactive apatitic crystals on natural polymers by reproducing the cascade of phenomena occurring in vivo during new bone formation, with the purpose of obtaining scaffolds which are able to trigger bone regeneration processes only by virtue of their compositional, morphological, and ultrastructural properties [79]. These bio-inspired mineralization processes involved the use of bio-polymers, particularly type-I collagen fibrils, which are dispersed in an aqueous solution of ions generally implicated in bone formation processes, the amount of which can determine the extent of the mineral phase in the bio-composite. An interesting aspect is that in these hybrid constructs the mineral phase is not simply embedded, but heterogeneously nucleated on a bio-organic matrix, thus more closely mimicking the ultrastructural features of bone tissue. In this context, fibrous hybrid materials could also be obtained as a multilayered device reproducing the different mineral content present in multifunctional tissues, such as osteochondral or periodontal regions [79][80][81][82]. This approach was intended to respond to a relevant clinical need related to the repair of osteochondral defects [83][84] using a 3D biomimetic scaffold that shows at the same time bioactive composition, porous fibrous structure, and good malleability, which is able to mimic the different osteochondral regions, namely bone, tidemark, and articular cartilage.
As natural polymers are easily subjected to fast enzymatic dissolution, the resorptive properties of hybrid scaffolds can be modulated using specific cross-linking additives, such as 1,4-butanediol diglycidyl ether (BDDGE), genipin or ribose, which are suitable for modulating the strength of interfibrillar bonds and in turn, the bioresorbability, hydrophilic properties, and mechanical performance of the bio-organic template [85][86][87]. In the last decade, the nucleation of apatite crystals in the presence of agar-gelatin hybrid hydrogel have been examined to understand the cooperative effects of the organic templates on the formation of the inorganic apatite phase, with favorable effects on the proliferation and differentiation of osteoblast-like MC3T3-E1 cells [88][89]. This system is suitable for the study of bio-inspired mineralization processes since (a) gelatin could be incorporated into the hydrogel and remain stable during the experiment, (b) the local concentration for crystallization is readily achievable, and (c) the deposits are easily harvested from the medium without agar contamination.
Bio-inspired hybrid materials possess excellent abilities in bone defect regeneration and also offer the possibility of developing new promising alternatives for the regeneration of osteochondral defects [90]. The regenerative ability boosted by high mimicry of osteochondral regions was attested by an in vivo study with collagen/apatite hybrids on sheep, showing that the various layers of the scaffold induced specific cell differentiation into osteoblasts (in the bony region) and chondrocytes (in the cartilaginous region) with the formation of osteochondral tissue with ordered histoarchitecture [91]. Hybrid collagen/apatite osteochondral scaffolds were also applied in various clinical trials. In most cases, the original functional structure of the cartilage was recovered within 2 years, as attested by numerous bio-markers that report the successful remodeling of the original fibrocartilage tissue into hyaline cartilage [92][93][94][95]. A noticeable aspect of the biomimetic approach, confirmed also by clinical studies, is the attainment of devices that are able to modulate the cell behavior only, thanks to physicochemical and structural features, without using any added growth factors. However, these hybrid, fibrous structures generally lack mechanical properties, enabling their application in load-bearing regions in order that technological advances in the incoming decades are highly desired and expected, in the perspective of reducing the recourse to invasive metallic joint reconstruction prosthesis, which is currently the only available solution for patients affected by osteochondral diseases.

3.3. Bioactive Glass Scaffolds

The development of silica glassy phases is among the first approaches considered to generate bioactive ceramics for bone regeneration. Bioactive glass structure is primarily composed of an interconnected open network of SiO4 tetrahedra that does not possess a uniform arrangement. This peculiar state favors the insertion of cations of various nature, referred to as network modifiers. The disruption of the O-Si-O network caused by the insertion of cations, such as Na+, K+, and Ca2+ and subsequent formation of non-bridging oxygen moieties result in high surface reactivity of these materials in an aqueous environment. Insertion of network modifiers, such as CaO and Na2O in a network comprising SiO2 and P2O5 is one of the most studied formulations, which gives rise to the formation of bioactive glasses with many applications in bone tissue engineering, as the crystallization of apatite-like phases is obtained on the surface of the glassy phase [96]. Magic-angle spinning nuclear magnetic resonance (MAS-NMR) spectroscopy has been applied for the study of the surface reaction of mesoporous bioactive glass (MBG) immersed in simulated body fluid (SBF) coupled with XRD analysis, proving the formation of amorphous calcium phosphate and subsequent crystallization into carbonated apatite within the first 4 to 24 h of SBF exposure, respectively. After 1 week, nanocrystalline carbonated apatite constitutes the main fraction (≈ 60%) of the total amount of phosphorus-bearing phases [97]. The ability of bioactive glasses to bond to bone tissue is a result of their chemical reactivity in physiological media and involves multiple steps:
  • Leaching through the exchange of protons from the physiological medium with labile network-modifying ions, such as Na+, K+, Ca2+, Mg2+, etc.:
Si − O − Na+ + H+ + OH− → Si − OH+ + Na+(solution) + OH−
The cation-exchange process increases the concentration of hydroxyl ions at the bioactive glass–solution interface, thereby raising the pH.
2.
The previous pH rise facilitates dissolution of the network and formation of additional silanol groups according to the reaction:
Si − O − Si + H2O → 2 [Si − OH+]
as well as the loss of soluble silica as Si(OH)4 passes into the solution.
3.
Polymerization of the SiO2− rich layer through condensation of neighboring Si–OH groups, which produces a layer rich in amorphous silica.
4.
Migration of Ca2+ ions to the surface of the silica-rich layer to form an amorphous film rich in CaO–P2O5, followed by thickening of the film by incorporation of soluble Ca2+ and PO43− ions from the solution.
5.
Crystallization of the amorphous CaO–P2O5 film by incorporation of OH− and CO32− from the solution, to form a carbonated apatite [98][99].
The controlled release of ions from bio-glasses in this context provides chemical signals and soluble Si and Ca ions are believed to be critical for controlling osteogenesis and/or activating genes responsible for osteogenesis in biological environment [100]. Studies have shown that MBGs do not inhibit osteoclastogenesis and allow for macrophage proliferation without inducing polarization toward M1 pro-inflammatory phenotype, indicating that MBGs would allow for the innate immune response required for the healing process without further inflammatory complications [101][102].
A major drawback in the development of bio-glass scaffolds is linked to the difficulty in assembling these materials into 3D scaffolds with appropriate mechanical properties. Indeed, upon heat treatment suitable for consolidation processes, bioactive glasses tend to undergo transformation into crystalline phases. Therefore, various attempts have been made to design fabrication strategies focused on the optimization of the glassy phase composition (for instance, by increasing the calcium/alkali ratio and partially replacing sodium with potassium) [103] to increase the crystallization temperature. In a different approach, glassy phases were used in the mixture with polymers to obtain reinforced composite materials [104][105][106]. As previously stated, the use of polymers in the design of biocompatible scaffolds has intrinsic problems related to degradation products, which are formed during bone resorption processes. However, the use of Bioglass® as an additive to PLGA foams has been reported to provide pH buffering effects at the polymer surface, which is assessed by long-term incubation tests (i.e., 30 days) [107][108], thus promising to reduce the drawbacks related to the use of polymers in bone scaffolds.

3.4. Self-Hardening Apatitic Scaffolds

When designing a 3D scaffold for bone substitution, it is necessary to consider the limitations imposed by the clinical procedure. Although most of the biomaterials designed to date have outstanding mechanical and biological properties after in vitro and in vivo trials, effective clinical applications are often limited by the difficulty in adapting the scaffold to complex shape defects or in accessing specific anatomical districts, such as for instance, the spine or the femur head or even by fixation problems [109]. Since in the area of regenerative medicine and dentistry CaPs have been already largely investigated due to their biocompatibility and osteoconductive behavior, new self-hardening materials, such as calcium phosphate cements (CPC) have been developed since decades, showing the ability to perfectly adapt to hard tissue defects (tooth, bone, etc.) [110].
Generally, CPC formulation involves the use of combination of calcium orthophosphates, which upon mixing with aqueous solutions forms a paste, which is able to harden after the implantation process at body temperature in the targeted area. Setting processes are primarily the result of dissolution processes involving the reagents and the precipitation of the final product. Classical CPC that employs α-tricalcium phosphate as the active reagent yields the formation of hydroxyapatite as a consequence of the hydrolysis reaction:
Alternatively, other types of CPC rely on acid-base reactions that entail the use of alkaline tetra calcium phosphate (TTCP) mixed with acidic di-calcium phosphate anhydrous (DCPA) following the reaction:
Other cement formulations provide different reaction products upon setting, such as brushite (CaHPO4·2H2O). As the hardening process occurs in a biological environment, the resulting apatite has strong similarities to biological apatite in terms of low crystallinity, high specific surface, bioactivity, and bioresorbability [19]. The peculiar morphology and hardening process rely on the physical interlocking of elongated HA crystals that form upon the hydrolysis process, thus resulting into micro/nanoporous architectures. What designates CPCs as versatile products are the plentiful variables that can be altered to deliver final products with specific rheological and mechanical properties. Main parameters affecting the porosity are, for example, the liquid/solid ratio and the particle size of the reagents used, that influence substantially the dissolution/precipitation processes [109][111]. The insertion of additives, such as biocompatible mannitol, for example, assisted the fabrication of bi-porous apatite without any inhibitory effects on the transformation into apatitic phases [112]. Although promising for bone regeneration purposes, it is widely accepted that CPCs suffer to satisfy clinical requirements, such as difficulty in obtaining good injectability and cohesion, as well as a general lack of mechanical strength, toughness, and brittleness, limiting their application in load-bearing defect sites [16][110]. The study of the CPCs scaffold evolution in biological environment highlights additional problems, such as the disintegration of the CPC paste upon early contact with biological fluids, such as blood due to weak cohesion and low porosity, which hinders homogeneous colonization of bone resorbing cells, compromising bioresorbability of the resulting scaffold, and limiting their extensive use in orthopaedics [110].

3.5. Mechanically Bearing, Biomorphic 3D Scaffolds

As previously stated, bone mechano-transduction process is a biologic phenomenon, which is able to effectively translate external mechanical stimuli exerted on the skeletal system into bio-electric signals, that can instruct bone cells to activate and sustain the continuous bone remodeling and self-repairing upon damage. Therefore, its activation is very important when it comes to regenerating the load-bearing bone parts. However, the reproduction of the complex biomechanical ability of bone tissue is quite a challenge. To date, it prevented the achievement of scaffolds which are effective in healing load-bearing bone defects [113].
A relevant example is the case of segmental bone defects (injuries in which a section of bone is completely shattered and/or absent), which the body cannot heal on its own. The common surgical procedures refer to the use of metallic components or bone bank pieces, which however cannot assure the regeneration of long bones with all its biomechanic functions and often results in adverse complications, such as pseudo-arthrosis, malunions, and loss of function [114][115][116].
A major limiting factor arises from insufficient vascularization, particularly when large bone defects have to be treated, which could ultimately lead to inefficient nutrient supply in the inner regions of the scaffold causing bone necrosis or insufficient cell penetration [60][117][118]. This drawback is the consequence of scaffolds with inappropriate composition favoring cell conduction, and/or insufficient pore structure and interconnection. In this context, novel approaches in scaffold development increasingly look at nature and involve the use of biostructures as sacrificial templates to fabricate advanced materials with morphological features replicating those of living organisms [119][120]. In this context, the great abundance of biologic structures with outstanding mechanical performance is a unique source of inspiration for material scientists for the production of implantable devices by structural replication of these natural sources.
Aragonite CaCO3-based crystals at the basis of the skeleton of common scleractinian, reef-building, colonial coral Porites (total porosity below 60 vol%, pore size range within 140–160 µm with all the pores interconnected), and Goniopora (total porosity above 70 vol%, large pore size ranging from 200 to 1000 µm) have been studied as a suitable template for the formation of biomorphic scaffolds with controlled porosity through the so-called “replamineform process” (meaning “replicated life forms”) [121]. The general synthesis of coral-derived biomorphic apatites (BA) involves the thermal treatment and decomposition of CaCO3 into CaO, followed by chemical exchange with (NH4)2HPO4 that leads to the conversion into HA, as the most thermodynamically stable CaP phase. Although this method allowed for the production of scaffolds with controlled micro-architecture, coral-derived HA implants developed to date were associated with brittleness, thus preventing effective surgical procedures. Indeed, it was shown that the unique mechanical performance of bone tissue is strongly related to its multi-scale hierarchic structure, which is lacking in most of the conventional ceramic devices, including coral-derived ones, as corals have a structure quite different from that of bone.

This entry is adapted from the peer-reviewed paper 10.3390/biomimetics7030112

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