**1. Introduction**

Musculoskeletal diseases are a worldwide cause of disability and pain, as they involve bones, teeth and joints, which are anatomical districts relevant for structural support, handling, protection, locomotion, mastication and many other physiological functions [1–3].

Bones are complex structures continuously undergoing dynamic remodeling due to a complex interaction of multiple biochemical processes, primarily ascribable on two different cell lines, osteoblasts and osteoclasts, as actors of bone deposition and resorption, respectively. Such processes can occur spontaneously in the case of minimal bone damage; however, if massive bone defects occur, as a result of a metabolic or traumatic cause, the physiological bone healing process has to be supported by a solid 3D scaffold, acting as a physical and instructive guide for cells [4–8].

Some properties requested for ideal bone scaffolds include *biocompatibility*, which is the ability of a biomaterial to function in vivo without eliciting any adverse side effects; *bioactivity*, which is the additional ability of a biomaterial to chemically bond with the surrounding tissue and to participate in specific biologically relevant phenomena (e.g., ion exchange); and *bioresorbability*, which is the ultimate ability of the implanted material to be resorbed over time, by active participation in physiological turnover reactions, favoring the formation of new tissue [9–12]. More specifically, scaffolds should exhibit *osteoinductivity* and *osteoconductivity*, both stimulating the osteointegration of the scaffold, which consists of a direct bone–scaffold interaction without fibrous tissue at the interface, essential to ensure mechanical stability and also the in-growth of blood vessels. In this respect, a leading concept guiding scaffold development is the achievement of high

**Citation:** Tavoni, M.; Dapporto, M.; Tampieri, A.; Sprio, S. Bioactive Calcium Phosphate-Based Composites for Bone Regeneration. *J. Compos. Sci.* **2021**, *5*, 227. https:// doi.org/10.3390/jcs5090227

Academic Editor: Francesco Tornabene

Received: 28 July 2021 Accepted: 18 August 2021 Published: 27 August 2021

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**Copyright:** © 2021 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

mimicry with targeted bony tissues, aiming to achieve a physiological cell response while preventing adverse foreign body reactions [13].

The design and development of biomimetic bone scaffolds have to be inspired by the complex physiological bone composition and structure. The bone microstructure is the result of the biomineralization of type I collagen, secreted by fibroblasts and osteoblasts cells, as a major component of the extracellular matrix of skin, tendon and bone [14]. Osteoblasts create the nano-composite structure of bone by secreting the ions responsible for the formation of apatite crystals. In turn, the ECM influences the adhesion, proliferation and differentiation of osteoblast, osteoclast and osteocyte [5,12]. The ECM is composed of inorganic and organic phases and water: the organic component consists of collagen and non-collagenous proteins, and the inorganic component contains calcium phosphate (mainly plate-like nanocrystalline hydroxyapatite, HA), calcium carbonate, magnesium phosphate and magnesium fluoride doped with various anionic (HPO4 <sup>2</sup>−, CO3 <sup>2</sup><sup>−</sup> and Cl−) and cationic species (Na+, K+, Mg2+, Sr2+, Zn2+, Ba2+, Cu2+, Al3+, Fe2+ and Si2+) trapped in the crystal structure. Carbonate ions are found in extent up to 8 wt%, while Na+, Mg2+, K+, Sr2+, Zn2+, Ba2+, Cu2+, Al3+, Fe2+/3+, F−, Cl<sup>−</sup> and Si4+ ions occur at trace (<1 wt%) [15]. Biogenic HA in bony tissue is non-stoichiometric with a Ca/P ratio between 1.5 and 1.67, where the inclusion of foreign ions in the crystal structure influences solubility, bioactivity, surface chemistry and morphology [16,17]. The general chemical formula for biogenic apatite is Ca10−x(PO4)6−x(HPO4 or CO3)x(OH or <sup>1</sup> <sup>2</sup> CO3)2−<sup>x</sup> with 0 ≤ x ≤ 2. One of the most common doping ions in biogenic HA is CO3 <sup>2</sup><sup>−</sup> ions, which can replace both phosphate and hydroxyl ions (leading to type B and type A carbonated apatite, respectively). For example, in B-type carbonated HA, the presence of CO3 <sup>2</sup><sup>−</sup> ions in the phosphate site inhibits the crystal growth and decreases the crystallinity; this structural disorder increases the chemical reactivity and enhances the solubility without changing the affinity of the osteoblast cells. Other possible anionic substitutions are with fluoride and chloride ions [17,18]. Cationic substitutions generally involve monovalent and bivalent cationic in the calcium sites of HA crystal lattice as reported in Table 1 [18–20].

**Table 1.** Relevant cation substitutions in natural HA crystal structure.


The bone structure exhibits a complex hierarchical architecture resulting from complex interactions of multilevel components, from micrometric osteons to apatite nanocrystals [21] (Table 2).

**Table 2.** Main components of bone structure, from macroscale to nanoscale.


In particular, it is possible to classify the levels and structures of components as follows:


Such a complexity is the main responsible factor for the outstanding mechanical performance of bone and its self-repair ability [23].

The ideal bone scaffolds should be endowed by several physico-chemical features, including chemical composition mimicking both the natural bone ECM and mineral phase, open and interconnected porosity capable of promoting neo-vascularization, tissue ingrowth, nutrient and oxygen supply, nano-structured surface topography positively driving adhesion, proliferation and differentiation of cells, that are adequate mechanical properties able to sustain the biomechanical loads toward the effective regeneration of the tissue.

Several studies have been carried out on the research of biomaterials, such as metals, natural or synthetic polymers, ceramics and composites, which can match all these characteristics, but no one fully satisfies all these requirements [24–30]. In particular, bioceramic-based scaffolds are widely used in numerous biomedical applications, including maxillofacial reconstruction, the stabilization of jaw bones, periodontal disease, as space fillers, self-hardening bone pastes/cements and as a coating on implants, due to their positive interaction with human tissue. Bioceramic-based materials can be classified as bioactive and bioinert materials. Ceramics considered as bioinert include alumina and zirconia; they show high chemical stability in vivo as well as high mechanical strength. However, they do not have osteogenic properties [31]. Bioactive ceramics, such as calcium phosphates (CaPs), silicates, bioactive glass, and titanium oxide, are capable of interacting with cells and thus able to promote and stimulate bone regeneration [28–33].

CaP bioceramics are widely used as bone substitutes since the 1920s and are considered as the golden standard in bone regeneration due to their similarity to the inorganic bone [34–37]. The chemical composition of CaPs relies on multiple ions, including calcium (Ca2+), orthophosphate (PO4 <sup>3</sup>−), metaphosphate (PO3 −), pyrophosphate (P2O7 <sup>4</sup>−) and hydroxide (OH−) [9,37] (Table 3).


**Table 3.** Some CaP materials: name, abbreviation, chemical formula, Ca/P ratio and solubility.

The solubility of CaP compounds strongly influences their behavior in vivo [37].

Among CaPs, HA is particularly promising for bone tissue regeneration due to its very close composition with natural apatite. In the last decades, the synthesis of HA has been investigated for different applications, including scaffolds, injectable pastes/cements, coatings for metallic implants and in nanomedicine as drug delivery platforms [38,39].

HA can be produced by several methods: high-temperature solid-state reactions or low-temperature precipitation [38]. Stoichiometric HA exhibits high stability at physiological pH, limiting its long-term resorption. Therefore, various recent studies have been focused on increasing the solubility and osteogenic activity of HA by ionic doping [39,40].

The notable interest in TCP comes from the combination of its solubility and low Ca/P ratio, particularly interesting when obtaining apatite crystals in an aqueous environment [16]. There are two polymorphs of TCP: the high-temperature α-TCP and the low-temperature β-TCP[41]. The β-TCP polymorph is stable at room temperature, while a transformation into α-TCP occurs at temperatures higher than 1125 ◦C. Besides a similar chemical composition, the TCP polymorphs have different crystalline structures, density and solubility, thus also resulting in different biological performance. The α-TCP phase is more soluble than β-TCP and can be easily hydrolyzed in calcium-deficient hydroxyapatite (1).

$$\text{3 Ca}\_3\text{(PO}\_4\text{)}\_2 + \text{H}\_2\text{O} \rightarrow \text{Ca}\_9\text{(HPO}\_4\text{)(PO}\_4\text{)}\_5\text{OH} \tag{1}$$

In addition, several ions can be introduced in the structure of TCP (Mg2+, Sr2+, Zn2+, Si2+, etc.), opening different thermodynamic scenarios in terms of polymorph stabilization; e.g., silicon was reported to stabilize α-TCP, while magnesium ions stabilize β-TCP.

Due to its high solubility, TCP has been used for the preparation of biphasic CaP scaffolds, able to conjugate the osteogenic properties of HA and the resorption behavior of TCP [42,43].

DCPD is biocompatible, biodegradable and osteoconductive [9]. DCPD can be prepared by the neutralization of phosphoric acid with calcium hydroxide at pH 3–4 at room temperature. DCPD can be obtained by double decomposition between calcium- and phosphatecontaining solutions in slightly acidic media. It can also be formed by the conversion of calcium phosphate salts, in acidic media, or by the reaction of calcium salts, such as calcium carbonate in acidic orthophosphate solutions. In vivo studies showed that DCPD converts into HA or it degrades and is replaced by bone [44–46]. Brushite, in medicine, is used in CaP paste/cement and as an intermediate for tooth remineralization [44,47].

Other silica-based bioceramics have also been studied as bone scaffolds, including wollastonite (CaSiO3), larnite (Ca2SiO4), hatrurite (Ca3SiO5), monticellite (CaMgSiO4), diopside (CaMgSi2O6), akermanite (Ca2MgSi2O7), merwinite (Ca3MgSi2O8), silicocarnotite (Ca5(PO4)2SiO4), nagelschmidtite (Ca7(SiO4)3(PO4)) and bioglass [48]. Silicon ions participate in bone metabolism, and silica-based materials exhibit good biological response in vitro, resulting in bioactive, biocompatible, bioresorbable, osteoinductive and osteoconductive behavior. The favored formation of apatite in physiological fluid was reported, thus facilitating the chemical interaction into the living bone structure following implantation [29,32].

The following steps explain the formation of apatite on the surface of silica-based bioceramics:


• The generation of ECM by the osteoblast to form new bone and its crystallization in the living composite structure.

Bioglasses are also a class of bioactive, osteoconductive and osteoinductive materials essentially composed of silicate, calcium, sodium and phosphate (e.g., composition of Bioglass 45S5® (wt%) 45 SiO2, 24.5 CaO, 24.5 Na2O, 6 P2O5). Upon implantation, bioglasses are not surrounded by fibrous tissue but form a strong, integrated bond to bone. In fact, when immersed in body fluids, the formation of a silica-rich layer on its surface takes place, which converts to a silica-CaO/P2O5-rich gel layer as a precursor of HA layer formation [24,33]. In addition, they are able to release ions, which enhance gene up-regulation and favor bio-degradation, in turn favoring bone regeneration [49]. Major drawbacks are related to the difficult consolidation of bioglasses into 3D porous scaffolds, as the required thermal treatment easily provokes the crystallization of oxides, thus losing the bioactive properties related to the material in its amorphous state. Therefore, alternative consolidation methods are currently under investigation; however, a major issue remains regarding the achievement of substantial mechanical properties associated with open porosity [50].

The mechanical properties of scaffolds play an important role in bone tissue engineering. The relevant mechanical properties of bone include Young's modulus, toughness, shear modulus, tensile strength, fatigue and compressive strength. Several approaches have been reported to increase the mechanical performance and load transfer efficiency between the scaffold and the surrounding bone tissue, mainly related to stronger interfacial bonding of the coating layer to the substrate [51].

The mechanical strength of ceramics mainly relies on their chemical composition, grain size, porosity extent and internal structural defects [37] (Table 4).

**Table 4.** Ideal features of scaffolds for bone regeneration, with respective proposed strategies to improve them.


Bioceramics typically exhibit higher compressive than tensile strength, but they are also intrinsically brittle, leading to sudden failure during handling and fixation [52]. In this respect, a critical challenge is related to the optimization of toughening mechanisms for ceramics [53,54].

The enhancement of the performance of bioceramic scaffolds has been widely explored by the combination of different calcium phosphate phases into bioceramic composites. The present work aims to provide the reader with an overview about the recently reported strategies to enhance the biofunctionality and mechanical properties of bioceramic scaffolds. In particular, various manufacturing techniques are explored, including the replica method, the sacrificial template, direct foaming, the low-temperature self-hardening method and biomorphic and biomimetic synthesis, as well as 3D printing, while also highlighting

future perspectives for the development of bioactive ceramic composites and devices with enhanced biofunctional properties (Figure 1).

**Figure 1.** Flow chart of biological and structural requirements underlying the ideal scaffold for bone tissue regeneration.

### **2. Fabrication of Bioceramic Composites**

The biological events occurring upon implantation of a scaffold for bone regeneration are strongly influenced by pore size distribution. The scaffold porosity affects the capability of the surrounding tissue to promote cell infiltration, migration, vascularization and nutrient and oxygen flows [18,55]. The morphological properties of scaffolds in terms of pore volume and size are important at both the macroscopic and the microscopic level.

It was reported that osteointegration and angiogenesis can be favored by interconnected macroporosity (100–600 μm) with channel-like microporosity [18]. A pore size increase is generally associated with an increase in permeability and the new bone ingrowth, while small pores are more suitable for soft tissue in-growth.

Over the past two decades, several technologies have been developed for the manufacturing of highly porous bioceramic-based scaffold for bone tissue regeneration [15,17,31,40–47]. In the next paragraphs, we explore the main fabrication techniques of porous scaffolds: traditional methods (partial sintering, replica method, sacrificial template and direct foaming), low-temperature self-hardening methods, biomorphic and biomimetic synthesis and 3D printing technology.

#### *2.1. Macroporous Compositescaffolds*

The development of materials with tailored porosity has been a matter of intense research in the last decades, particularly in the case of composite scaffolds for bone tissue regeneration, because of the crucial role of voids in the structure to guide and facilitate cell proliferation and neovascularization [56].

One of the first reported approaches to tune the porosity of ceramics was the partial sintering process: the pore size distribution is mainly affected by powder particle size and sintering temperature, as higher sintering temperatures induce a significant decrease in intergranular porosity [57,58].

A great research effort has also been devoted to the preparation of macroporous bioceramic scaffolds, leading to the establishment of various techniques, including template-assisted (replica and sacrificial template) and template-free techniques (direct foaming) [56,59,60] (Table 5) [56,60].


**Table 5.** Main processing steps involved in the fabrication of porous bioceramics.

These methods generally involve the preparation of slurries, intended as aqueous suspensions of dispersed powders; then, the slurries are properly manipulated, dried and thermally consolidated.

The replica method is a template-assisted technique based on the impregnation of a polymeric sponge with a defined porous structure and pore size into the ceramic slurry in order to produce microporous structures exhibiting the original sponge morphology [56]. The templates used in this technique can be either synthetic or natural polymers (e.g., polyurethane and cellulose, respectively). The macroporous scaffolds obtained with this method can reach an anisotropic porosity ranging from 40 to 95% and are characterized by a cross-linked structure with highly interconnected pores ranging in size from 200 μm to 3 mm [56].

The sacrificial template method involves the homogeneous dispersion of sacrificial phases into a continuous matrix of ceramic particles or ceramic precursors, followed by drying and sintering. A wide variety of sacrificial materials can be used as pore-forming agents, including natural polymers (e.g., gelatin, potato starch, cotton), synthetic polymers (e.g., polymer beads, organic fibers, polyethylene) and inorganic polymers (e.g., NaCl, K2SO4). The removal of sacrificial materials from the matrix can be achieved by thermal treatments or chemical processes. This method leads to porosity ranging from 20 to 90%, with an average pore diameter of 1–700 μm [18,56].

Template-free foaming techniques are particularly promising due to the absence of massive amounts of organic phases to be eliminated during thermal consolidation. Direct foaming represents an easy, cheap and fast way to prepare macroporous bioceramics with open porosity from 40 to 97% and pore size 10 μm–1 mm by incorporating gas bubbles into ceramic slurries, followed by drying and sintering [18,56,61]. The total porosity volume is related to the amount of gas bubbles incorporated during the foaming process, whereas the pore size depends on the stability of the poured foam before drying [18,56,61].

The sacrificial template approach also includes the freeze-casting method, which is based on the controlled freezing of liquid-based ceramic slurries [18]. The freezing of the liquid, generally water, induces the formation of anisotropic ice structures, intended as fugitive materials, during the subsequent freeze-drying process [62]. The efficacy of the process is affected by several parameters, including the viscosity of the slurry, the solvent and the freezing control in space and time. Typical structures obtained by freeze-casting methods showed well-defined pore connectivity along with directional and completely open porosity, such as a lamellar morphology after sintering [63]. The channellike anisotropic porosity obtained by the freeze-casting method may lead to scaffolds with channels similar to cortical bone, particularly useful for long bone applications [18].

#### *2.2. Self-Hardening Bioceramic Composites*

The possibility to obtain bioactive ceramics through low-temperature self-hardening processes has been widely explored in the form of bone cements for injectable orthopedic applications, including spinal fusion, vertebroplasty and kyphoplasty [30,64–66]. Bone cements refer to pastes able to self-harden under physiological conditions and can be injected in vivo through minimally invasive surgery [64]. The first bone cement used in orthopedics was based on polymers, in particular polymethylmethacrylate (PMMA) in 1958, and, in the 1970s, the FDA approved bone cement for use in hip and knee prosthetic fixation [67]. Despite PMMA-based cements exhibiting good handling, setting times and mechanical performance, they are not osteogenic nor bioresorbable. Calcium phosphate cements (CPC) were discovered by Brown and Chow in the 1980s [68–70], overcoming the drawbacks of PMMA cements in terms of exothermic polymerization hardening and chemical composition. In this respect, CPCs exhibit bioactivity, bioresorbability and a physiological hardening at 37 ◦C, also allowing the incorporation of biomolecules [68]. The main drawback of CPCs hampering their clinical applications is related to their poor mechanical performance, which limits their applicability to a moderate- or non-loadbearing situation [71].

CPCs can be classified by several parameters, including the number of components in the solid phase, the type of setting reaction and the type of end product (Table 6) [38,68].


**Table 6.** Classification of CPC.

Many different formulations of CPCs have been developed, and they can be divided into two groups based on the type of end product: brushite (DCPD) and apatite (HA or CDHA) cements. Both brushite and apatite CPCs are produced upon mixing one or more CaP powders with aqueous solutions, which induces the dissolution of the initial CaPs; this is followed by precipitation into crystals of DCPD, HA or CDHA depending on the compositions of the powders and the setting reactions that take place [38,72]. During precipitation, new apatitic crystals grow and their physical entanglement causes the hardening or setting at body temperature.

Apatitic CPCs can be obtained by mixing single or multi-components with aqueous solutions that undergo hydrolysis or acid–base reactions, respectively. In the first case, the end product is calcium-deficient hydroxyapatite (CDHA), and in the latter, it is stoichiometric HA [64,68]. Some examples are as follows:


$$\text{C}\ \text{a-Ca}\_3\text{(PO}\_4\text{)}\_2 + \text{H}\_2\text{O} \rightarrow \text{Ca}\_9\text{(HPO}\_4\text{)(PO}\_4\text{)}\_5\text{(OH)}\_2$$


$$\text{Ca}\_4(\text{PO}\_4)\_2\text{O} + \text{CaHPO}\_4 \rightarrow \text{Ca}\_5(\text{PO}\_4)\_3(\text{OH})\_2$$

Brushite CPC obtained by an acid–base reaction between TCP (almost neutral) and monocalcium phosphate monohydrate, MCPM (acidic):

$$\text{\{\\$-Ca\_3(PO\_4)\_2 + Ca(H\_2PO\_4)\_2\cdot H\_2O + 7H\_2O \rightarrow 4CaHPO\_4\cdot 2H\_2O}}$$

Two of the most important parameters that play a key role in the final CPC features are the liquid-to-powder ratio (LPR) and the particle size of the starting powder [37,68]. The LPR influences setting time, injectability, cohesion, mechanical properties and the porosity of harder CPC [73]. The setting time is the "time required from the start of powdered agent and liquid agent blending until hardening of the cement", according to ISO/DIS 18531 for CaPs [30,74], and influences the clinical applicability of both apatite and brushite cements as well as their injectability [30,74].

Both particle size and the LPR influence the final surface morphology of the brushite or apatite crystals and the total porosity of the final scaffolds, which affects the mechanical performance and the resorbability of scaffolds and therefore the overall bioactivity (Table 7) [37,68]. The reduction in the particle size of CaPs increases the surface area, thus affecting the reaction kinetics and yielding small needle-like crystals rather than large plate-like crystals as observed when larger CaP precursor particles are used [38,75]. Moreover, porosity is also attributed to the amount of liquid phase used; thus, by increasing the LPR, the amount of liquid phase decreases, and the porosity increases. This effect of the LPR explains the difference between brushite and apatite cement in terms of microstructure porosity: the water consumption during the setting reaction of brushite cement is larger than that of the apatite, which leads to the formation of a larger crystal size and makes the total porosity smaller and average pore size greater than those of the apatitic cements [37,73]. The typical porosity of CPC ranges between nano- and sub-micrometer size, allowing the flow of physiological fluids within the microstructure of the cement, but the pores are too small to facilitate the growth of bone tissue; in this regard, porogens are often used [69].

**Table 7.** Effect of particle size and liquid-to-powder ratio on the crystals' morphology and pore distribution.


As mentioned above, increasing porosity leads to decreasing mechanical strength; thus, a compromise must be sought between mechanical performance and porosity degree.

One of the advantages of CPC is the room-temperature self-hardening mechanism, which, combined with the intrinsic porosity, allows the incorporation of drugs, biologically active molecules and cells, obtaining drug delivery materials [76,77]. The incorporation of active molecules in CPCs can be achieved by dissolving it in the liquid phase or by a combination with the powder phase of the CPC mixing setting [68,78]. Another possible approach is the superficial adsorption of drugs on the CPC surface by incubation of the scaffold in the drug solution: the kinetic release of drugs depends on the functionalization, microstructure and resorbability of the CPC matrix [68,78].
