Calcium Orthophosphate (CaPO₄) Containing Composites for Biomedical Applications: Formulations, Properties, and Applications

Abstract

The goal of this review is to present a wide range of hybrid formulations and composites containing calcium orthophosphates (abbreviated as CaPO₄) that are suitable for use in biomedical applications and currently on the market. The bioactive, biocompatible, and osteoconductive properties of various CaPO₄-based formulations make them valuable in the rapidly developing field of biomedical research, both in vitro and in vivo. Due to the brittleness of CaPO₄, it is essential to combine the desired osteologic properties of ceramic CaPO₄ with those of other compounds to create novel, multifunctional bone graft biomaterials. Consequently, this analysis offers a thorough overview of the hybrid formulations and CaPO₄-based composites that are currently known. To do this, a comprehensive search of the literature on the subject was carried out in all significant databases to extract pertinent papers. There have been many formulations found with different material compositions, production methods, structural and bioactive features, and in vitro and in vivo properties. When these formulations contain additional biofunctional ingredients, such as drugs, proteins, enzymes, or antibacterial agents, they offer improved biomedical applications. Moreover, a lot of these formulations allow cell loading and promote the development of smart formulations based on CaPO₄. This evaluation also discusses basic problems and scientific difficulties that call for more investigation and advancements. It also indicates perspectives for the future.

1. Introduction

Bone fractures resulting from traumatic injury or age-related conditions are a common form of tissue damage. Orthopedic surgeons face significant challenges in surgically treating such fractures, particularly when dealing with large bone defects that require the implantation of temporary or permanent prostheses. The rapid aging of the population and the limitations of natural bone grafts only exacerbate the situation, leading to a high demand for bone replacement materials. However, the use of xenografts in medicine, such as bovine bone, can increase the risk of viral infection. Additionally, xenografts have a lower potential for bone formation, are highly immunogenic, and generally resorb more quickly than autogenous bone. Similar limitations exist for human allogeneic grafts, which are tissue transplants between individuals who are homologous but not genetically identical. The concerns about the risk of tumor cells and bacterial and viral infections, as well as potential immunological and blood group incompatibility, are even greater. Additionally, the collection and preservation of allogeneic grafts (exogenous bone) present additional limiting factors [1,2,3]. Autologous bone (also known as endogenous bone) remains the preferred choice among all other alternatives, owing to its exquisite osteoconductive, osteogenic, osteoinductive, and biocompatible features, as well as its non-allergic and non-toxic characteristics. Furthermore, it has bone matrix proteins and live osteogenic cells that encourage osteogenesis. The body generally accepts autografts and they quickly integrate with the adjacent bone tissue. Therefore, these devices are regularly utilized for prolonged periods of time, yielding favorable clinical results [2,3,4,5]. However, it is worth noting that there have been frequent instances of complications in the past [6]. Furthermore, the availability of autografts from the iliac crest or other areas of the patient’s body is regrettably restricted due to the limited number of donor sites. This may necessitate further healing of the donor site and may cause prolonged post-operative pain. Moreover, the utilization of biological transplantation in medicine results in additional trauma and scarring due to the removal of donor tissue during extra surgical procedures. Additionally, the limited amount of donor tissue, morbidity of the donor site, and potential risk of immunological incompatibility and disease transfer make biological transplantation an imperfect solution [7,8,9]. In this context, engineered materials (synthetic or alloplastic bone grafts) are a viable option due to their widespread availability and ability to be processed and tailored to meet specific needs. Furthermore, they are not associated with concerns such as infection, immunologic incompatibility, sterility, or donor-site morbidity. The research on developing engineered materials for bone tissue repair is a crucial challenge in the field of clinical biomaterials [10,11].

Several categories of synthetic bone graft biomaterials are suitable for in vivo use. They can be classified according to their source (biological, non-biological), materials (metals, ceramics, polymers, composites), material consistency (implantable solids, injectables, adhesives), and the presence or absence of porosity, living cells, nanoparticles, and other factors [12]. Natural corals, coral-derived materials, bovine porous demineralized bones, human demineralized bone matrix, bioactive glass, glass ceramics, and calcium orthophosphates (abbreviated as CaPO₄) are all materials used in bone regeneration [13]. Porous bioceramics composed of CaPO₄ are especially prominent due to their excellent biocompatibility and ability to bind to biological bones because CaPO₄ is a key component of mammalian calcified tissues, including teeth and bones [14,15]. A variety of CaPO₄-based biomaterials with diverse chemical structures are currently available for purchase on the market [16,17]. However, CaPO₄ alone does not possess the mechanical or elastic properties of natural calcified tissues. In other words, CaPO₄ scaffolds pose several concerns for mechanical performance after implantation due to their low elasticity, high brittleness, poor tensile strength, low reliability, and fracture toughness. Additionally, molding CaPO₄ into the desired shape proves difficult in many instances [16,17].

Partial flexibility coupled with the superior strength of natural biomineralized tissues (bones and teeth) are attributed to the presence of bioorganic polymers (predominantly collagen type I fibers) coupled with the natural ceramic (largely a poorly crystalline ion-substituted calcium-deficient hydroxyapatite (CDHA) phase, referred to as ‘bioapatite’)—refer to

Table 1. The biochemical composition * of bones [22].

Inorganic Phases	wt%	Bioorganic Phases	wt%
CaPO4 (biological apatite)	~60	collagen type I	~20
water	~9	non-collagenous proteins: osteocalcin, osteonectin, osteopontin, thrombospondin, morphogenetic proteins, sialoprotein, serum proteins	~3
carbonates	~4	other traces: polysaccharides, lipids, cytokines	balance
citrates	~0.9	primary bone cells: osteoblasts, osteocytes, osteoclasts	balance
sodium	~0.7
magnesium	~0.5
other traces: Cl−, F−, K+ Sr2+, Pb2+, Zn2+, Cu2+, Fe2+	balance

* The composition is varied from species to species and from bone to bone.

[18,19,20,21,22]. In bones, elastic collagen fibers align with the primary direction of tension. Demineralized bone has high flexibility and is easily bent, whereas collagen-free bone is highly brittle. This is due to inorganic nanoscale bioapatite crystals that confer stiffness and rigidity, whereas bioorganic fibers provide elasticity and toughness. In bones, the integration of both types of materials occurs at the nanometer scale, with the size of the crystallites, the orientation of the fibers, and the order within the components determining the nanostructure and, consequently, the function and mechanical features of the entire composite. From a mechanical standpoint, bone is highly sturdy at low strain rates, but exhibits brittleness at high strain rates. Additionally, bone is an anisotropic material due to the directional dependence of its properties [18,19,20,21].

Designing an optimal bone graft that mimics the structure and function of natural bone is a significant obstacle. Understanding bone structure is crucial before attempting to create a bone graft. According to conventional expectations, ideal bone grafts must possess certain characteristics. They should be benign, available in varied shapes and sizes, and possess adequate mechanical properties for use in load-bearing regions. They should be capable of forming chemical bonds at the bone–implant interface, while also being osteogenic, osteoinductive, osteoconductive, biocompatible, and fully biodegradable, to favor bone growth and moldable for filling and repairing bone defects [23,24]. Furthermore, ideal implants should have a chemical composition similar to bone, wherein the presence of CaPO₄ is essential. They should also exhibit continuous porosity, facilitating penetration of living host tissues, and have viscoelastic and semi-brittle behavior similar to bone [25,26,27]. Additionally, the desirable degradation rate of implants needs to match the healing rate of human tissues and be free from chemical and biological irritation and toxicity resulting from corrosion and degradation. Ideally, the mechanical strength of the implant and the transplanted bone should remain constant during the regeneration process [28]. Some argue that the mechanical properties of implants should exceed those of bones; therefore, in cases of severe trauma, it may be necessary to destroy the bone rather than the implant [23]. To enable clinical applications, grafts must be sterile, easily preservable, processable, and cost-effective. Unfortunately, there are currently no engineered biomaterials that fully meet these requirements and it seems unlikely that such materials will be developed in the near future. Most of the available bone grafting biomaterials are either osteogenic, osteoinductive, or osteoconductive [1].

The design of bone replacement materials necessitates thoughtful contemplation of the bone type and its mechanical properties. Specifically, bones that are highly loaded, such as the femur, require an implant with sufficient stiffness to provide stability, yet not so rigid as to cause strain retention. On the other hand, in applications with relatively low loads, like skull repair, having the appropriate three-dimensional shape is crucial for stability and aesthetic purposes. One promising option is to utilize materials with a composition and nanostructure similar to that of bone tissues. Organic–inorganic hybrid biomaterials development, which imitates the structure of calcified tissue and combats the limitations of individual materials, has great potential for enhancing conventional bone implants. Biologically important CaPO₄ and bioabsorbable polymers can be combined to create biocomposites with unique physical, biological, and mechanical properties and predictable degradation behavior [29]. The general characteristics of these biocomposites depend on the nature, structure, and relative content of the components, while other factors such as preparation conditions also play a role in determining the properties of the final material. Currently, CaPO₄ is applied as fillers or coatings, either in the form of particles or fibers, incorporated into or applied on top of biodegradable polymer matrices. Due to its improved physical, biological, and mechanical properties, it is slowly being considered as a scaffold for bone tissue engineering [30,31,32,33]. Furthermore, these biocomposites meet the overall criteria for next-generation biomaterials by combining bioactivity and bioabsorbability, activating in vivo tissue regeneration mechanisms, stimulating the body’s self-healing capacity, and resulting in implant replacement with regenerated tissue. Thus, an optimal material can be designed by effectively combining a ductile polymer matrix with a tough, bioactive particulate bioceramic filler. Ideally, this approach will yield superior structures suitable for use as implants and posterior dental restorations [29,34].

In clinical orthopedic practice, the initial composite materials were lint-reinforced plasters utilized by Mathijsen as external fixatives (bandages) in fracture treatment back in 1852 [35], followed by Dreesman in 1892 [36]. Subsequently, considerable advancements have been made in the clinical usage of dissimilar types of composite materials. It is possible to create and use a variety of composite materials with distinct mechanical and biological properties to satisfy a range of clinical needs [37]. Therefore, the purpose of this review is to outline the wide range of hybrid formulations and CaPO₄-based composites that are now available and appropriate for use in biomedical applications.

2. General Knowledge and Experience

According to Wikipedia, “A composite material (also called a composition material or shortened to composite, which is the common name) is a material which is produced from two or more constituent materials. These constituent materials have notably dissimilar chemical or physical properties and are merged to create a material with properties unlike those of the individual elements. Within the finished structure, the individual elements remain separate and distinct, distinguishing composites from mixtures and solid solutions” [38]. Composite materials are inherently heterogeneous, with each phase retaining its unique identity and properties while being bonded to maintain an interface. This results in enhanced specific or synergistic properties that cannot be achieved from the original phase alone [39]. In line with previous research, we also explore the following: “for the purpose of this review, composites are defined as those having a distinct phase distributed through their bulk, as opposed to modular or coated components” ([40], p. 1329). Therefore, with a few exceptions, we did not consider CaPO₄ deposits produced on different materials by various deposition techniques [41,42,43,44,45,46] or CaPO₄ coated with other compounds [47,48,49,50,51]. However, we did include composite coatings. Structures such as porous CaPO₄ scaffolds with cells filling the pores [52,53,54] and CaPO₄ impregnated with bioactive substances [55,56] can also be described as composites or hybrids, but we did not consider them in this study. Finally, biohybrids are also defined as “the functional combination of proteins, viable cells or microorganisms with non-biological materials” [57]; such compositions are also excluded.

Composite materials consist of two primary components: a matrix (also referred to as the continuous phase) and dispersed phase(s). The presence of at least one component from each category is essential to form a composite material. Typically, the matrix is considered to be the component constituting the major and continuous phase (making up >50% by volume, often exhibiting relatively low stiffness and strength), whereas the reinforcement is the one present as a minor, discontinuous, and/or dispersed phase (comprising <50% by volume, frequently showcasing relatively high stiffness and strength). An interface serves as a boundary between the constituents, typically located in a small area where the chemical composition significantly changes and forms a bond with one another, thereby playing a vital role in load transfer. The advantage of composites lies in their customizable properties, achieved by modifying ratios and placement of constituents, as well as adjusting interfaces. The literature contains a comprehensive overview of the primary manufacturing and processing techniques [40,58]. The matrix not only occupies the space but also upholds the dispersed phase by enveloping it and preserving its respective positions. One or several dispersed phases are usually responsible for augmenting one or several matrix properties. The main goal of the majority of composite materials is to improve the matrix’s mechanical properties, like strength and stiffness. To achieve the intended results, however, additional characteristics including biocompatibility, radiopacity, density, transport qualities (thermal or electrical), and erosion stability are also essential. This combined effect generates properties that are inexistent in individual component materials [58,59]. Control of the volume fraction and arrangement of dispersed phases allows for modification and adaptation of the properties and design of composite materials to meet specific conditions. For instance, in ceramics, the dispersed phase functions as a crack inhibitor and reinforcing material. To do this, techniques include bridging the crack face, deflecting the crack tip, absorbing energy during shrinkage, and producing stress redistribution in the vicinity of the crack tip [60]. The dispersed phase’s volume proportion, size, form, and orientation; the condition of the reinforcement/matrix contact; and the homogeneity of the composite material as a whole are additional variables to take into account with composites. A higher volume fraction of the reinforcement phase enhances the mechanical properties of composite materials. Continuous and aligned fibers are effective in preventing crack propagation by having an anisotropic behavior. Composites are structurally anisotropic, and their mechanical properties vary with orientation. Furthermore, in addition to functionally graded materials, a uniform distribution of dispersed phases is also desirable in order to provide composite materials with consistent properties [38,58,59].

Composites can be classified into four types: simple, complex, graded, and hierarchical. A simple composite contains one dispersed phase that is homogeneously distributed throughout the matrix. A complex composite, on the other hand, contains multiple dispersed phases that are homogeneously distributed within a matrix. A graded composite is intentionally structured in a heterogeneous manner by distributing one or more dispersed phases throughout the matrix. Hierarchical composites are structures whereby fine entities from a simple or complex composite combine to create coarse particles or granules. These particles are subsequently dispersed within another matrix, generating composite structures at a second hierarchical scale. Another classification categorizes composites into four groups: (i) fiber-reinforced composites, where fibers are embedded in the matrix; (ii) layered composites, where layers of different materials are incorporated to form a structure; (iii) particulate composites, where particles or flakes are dispersed in the matrix; and (iv) hybrid composites, which are combinations of any of the above types. Particulate composites typically employ micro- or nanoscale reinforcements made up of particles. In contrast, fibrous composites may consist of long, continuous fibers that are either aligned or woven and short fibers that are either aligned or randomly oriented. Another means of classifying composites is based on the matrix material. For example, there are ceramic, polymer, and metal composites [37].

In the design of composite materials, three interdependent factors must typically be taken into account: (i) the selection of suitable matrix and dispersion materials; (ii) the selection of appropriate manufacturing and processing methods; and (iii) the internal and external device design [40]. Additionally, all composite materials require shaping, which may involve adding a matrix material before or after placing the dispersion material on the mold surface or in the mold cavity. Matrix materials go through melting processes that can happen in various ways, like chemical polymerization, curing, or solidification from the molten state. The particular melting mechanism depends on the nature of the matrix material. Because of the overall heterogeneity, many composite materials exhibit orthotropic physical properties, meaning that different strengths or properties exist in different orthogonal directions [38,58,59].

Since composites require the mixing of two or more materials, the process of phase-mixing is essential [61,62]. Moreover, interfacial strength between the phases is crucial since inadequate adhesion between them can cause early degradation of the interface, leading to lower mechanical properties, particularly the tensile strength. From a chemical standpoint, various types of interactions exist among composite materials. These include materials with strong interactions (covalent, coordination, and ionic bonds), those with weak interactions (van der Waals forces, hydrogen bonds, hydrophilic–hydrophobic equilibrium), and materials with no chemical interactions between the components [63]. Furthermore, wetting plays a significant role in the bonding and adhesion of materials. The result is determined by the polar groups in the matrix and the hydrophilicity or polarity of the filler.

Biocomposites have two definitions. Firstly, they refer to a composite material that is safe and contains one or more components that can interact favorably with the human body in vivo, thus promoting healing processes and implant uptake [64]. Secondly, they are a type of biomaterial with composite properties. The most recent definition of a biomaterial is as follows: ‘A biomaterial is a material designed to take a form that can guide the process of treatment or diagnosis through its interaction with a living system’ [65]. In any case, the biocompatibility of biocomposites appears to be more crucial than other forms of compatibility [37,66,67].

Historically, according to the Scopus database, the earliest article featuring the term ‘biocomposite’ in the title was published in 1987 [68], while an article combining ‘biocomposite’ and CaPO₄ (it was hydroxyapatite, HA) in the title was published in 1991 [69]. However, a study published in 1976 [70] marks the earliest paper that combines the terms “composite” and HA in its title without the prefix “bio”. As of May 2024, Scopus reports 7633 articles with titles featuring the combination of “composite” and “apatite” and 2160 articles featuring “composite” and “calcium phosphate”. These numbers indicate an active field of research.

The most common properties of bio-organic and inorganic domains formulated into biocomposites are summarized in

Table 2. General respective properties from the bioorganic and inorganic domains to be combined in various composites and hybrid materials [24].

Inorganic	Bioorganic
hardness, brittleness	elasticity, plasticity
high density	low density
thermal stability	permeability
hydrophilicity	hydrophobicity
high refractive index	selective complexation
mixed valence slate (red-ox)	chemical reactivity
strength	bioactivity

[24], while, for an overview of the general advantages of modern CaPO₄-based biocomposites compared to CaPO₄ bioceramics and bioabsorbable polymers, readers are referred to “Composite materials strategy” section of Ref. [29].

3. Main Components of Bone Graft Biocomposites and Hybrid Biomaterials

3.1. CaPO₄

CaPO₄ was first identified as a significant constituent of bone in 1769 and has become the subject of constant study ever since [71,72]. One of CaPO₄’s primary applications as a material for bone replacement is its chemical resemblance to the mineral elements of mammalian teeth and bones [14,15,16]. CaPO₄ is biocompatible and non-toxic and exhibits bioactive behavior while assimilating into biological tissue via the same processes present in healthy bone remodeling. This results in a close physico-chemical bond between the CaPO₄ implant and bone, referred to as osteointegration. Moreover, CaPO₄ promotes osteoblast adhesion and proliferation, making it an essential component in various orthopedic applications. However, the mechanical properties of using CaPO₄ alone as a load-bearing biomaterial, such as brittleness, zero elasticity, and low fatigue resistance [23], limit its effectiveness. In porous ceramics and scaffolds, this poor mechanical behavior is more prominent, as voids larger than 100 µm are crucial for proper vascularization and osteocyte colony formation [73,74]. Therefore, in biomedical applications, CaPO₄ alone is used mainly as a filler or coating [16].

Table 3. Existing forms of CaPO₄ and their major properties [16].

Ca/P Molar Ratio	Compound	Formula	Solubility at 25 °C, −log(Ks)	Solubility at 25 °C, g/L	pH Stability Range in Aqueous Solutions at 25 °C
0.5	Monocalcium phosphate monohydrate (MCPM)	Ca(H2PO4)2·H2O	1.14	~18	0.0–2.0
0.5	Monocalcium phosphate anhydrous (MCPA or MCP)	Ca(H2PO4)2	1.14	~17	[c]
1.0	Dicalcium phosphate dihydrate (DCPD), mineral brushite	CaHPO4·2H2O	6.59	~0.088	2.0–6.0
1.0	Dicalcium phosphate anhydrous (DCPA or DCP), mineral monetite	CaHPO4	6.90	~0.048	[c]
1.33	Octacalcium phosphate (OCP)	Ca8(HPO4)2(PO4)4·5H2O	96.6	~0.0081	5.5–7.0
1.5	α-Tricalcium phosphate (α-TCP)	α-Ca3(PO4)2	25.5	~0.0025	[a]
1.5	β-Tricalcium phosphate (β-TCP)	β-Ca3(PO4)2	28.9	~0.0005	[a]
1.2–2.2	Amorphous calcium phosphates (ACP)	CaxHy(PO4)z·nH2O, n = 3–4.5; 15–20% H2O	[b]	[b]	~5–12 [d]
1.5–1.67	Calcium-deficient hydroxyapatite (CDHA or Ca-def HA) [e]	Ca10−x(HPO4)x(PO4)6−x(OH)2−x (0 < x < 1)	~85	~0.0094	6.5–9.5
1.67	Hydroxyapatite (HA, HAp or OHAp)	Ca10(PO4)6(OH)2	116.8	~0.0003	9.5–12
1.67	Fluorapatite (FA or FAp)	Ca10(PO4)6F2	120.0	~0.0002	7–12
1.67	Oxyapatite (OA, OAp or OXA) [f], mineral voelckerite	Ca10(PO4)6O	~69	~0.087	[a]
2.0	Tetracalcium phosphate (TTCP or TetCP), mineral hilgenstockite	Ca4(PO4)2O	38–44	~0.0007	[a]

^[a] These compounds cannot be precipitated from aqueous solutions. ^[b] Cannot be measured precisely. However, the following values were found: 25.7 ± 0.1 (pH = 7.40), 29.9 ± 0.1 (pH = 6.00), and 32.7 ± 0.1 (pH = 5.28). The comparative extent of dissolution in an acidic buffer is ACP >> α-TCP >> β-TCP > CDHA >> HA > FA. ^[c] Stable at temperatures above 100 °C. ^[d] Always metastable. ^[e] Occasionally it is called “precipitated HA (PHA)”. ^[f] Existence of OA remains questionable.

provides a comprehensive list of all known forms of CaPO₄, together with standard abbreviations and important features. Books and monographs on the subject can provide further in-depth information on CaPO₄ [16,75,76,77,78].

3.2. Polymers

Polymers are (bio)organic compounds consisting of big molecules covalently bound to thousands of smaller units (monomers) to form lengthy chains, resulting in highly malleable materials. In this regard, synthetic polymers can be likened to natural ones (lipids, proteins, and polysaccharides), which are the primary functional constituents of the biological environment. Polymers vary from one another in terms of their chemical composition, molecular weight, polydispersity, crystallinity, hydrophobicity, solubility, and thermal transition. Furthermore, through the copolymerization or blending of two or more polymers, as well as by adjusting the polymer type and chain length, their properties can be precisely altered throughout a wide range [79,80]. Due to their strong viscoelastic characteristics, polymers can be easily shaped into intricate porous networks and channels, spongy sheets, and gels [81]. Non-magnetic and X-ray transparent polymeric materials are completely interoperable with modern diagnostic techniques such as magnetic resonance imaging and computed tomography. Regrettably, the in vivo physiological environment presents stringent requirements, which many of these materials fail to meet. The primary criteria for polymers that are appropriate for biomedical applications are biocompatibility, lack of excessive or chronic inflammatory reactions after implantation, and degradation into non-toxic products. However, for load-bearing applications, most polymers lack the necessary stiffness, ductility, and ultimate mechanical properties. Thus, in the case of good biocompatibility, most polymeric materials are primarily employed for soft tissue replacement (e.g., skin, vascular, cartilage, and ligament replacements). Moreover, the characteristics of polymers can be impacted by sterilizing procedures, such as autoclaving, ethylene oxide, and ⁶⁰Co irradiation [82].

Several biocompatible polymers are suitable for biomedical applications [83,84,85]. For instance, researchers have investigated polyacrylates, poly(acrylonitrile-co-vinyl chloride), and polylysine for cell encapsulation and immunoseparation [86,87]. Additionally, researchers have examined polyorthoesters and PCL as drug delivery devices, with the latter’s slow degradation rate enabling long-term sustained release [88]. PCL is a semi-crystalline aliphatic linear polyester polymer with a low degradation rate of more than 2 years. It is both bioabsorbable and biocompatible, with an adequate absorption time and non-toxic by-products generated upon degradation and release. As a result, it finds extensive use in pharmaceuticals and wound dressings [89,90]. Likewise, PU is employed in hard and soft tissue engineering as well as nanomedicine [91]. Polyanhydrides are utilized as polymers for orthopedic purposes and can undergo photopolymerization in situ. Additionally, they are currently being researched as delivery devices for bone replacement and augmentation due to their rapid and clear surface erosion properties [88,92,93]. To counteract their suboptimal mechanical properties, researchers have either copolymerized the materials with imides or formulated them to be cross-linkable in situ [93]. Other polymers, such as polyphosphazene, are potential materials for regenerating skeletal tissue due to their modifiable properties (e.g., degradation rate) through changes in their side chain structure and their demonstrated ability to enhance osteoblast adhesion [93]. PPF exhibits outstanding mechanical properties (comparable to trabecular bone), can cross-link in vivo through C=C bonds, and has risen to prominence as a replacement material for bones due to its hydrolyzability. It is also considered a material for drug delivery devices [88,92,93,94,95,96,97,98,99]. Polycarbonate has been suggested as an appropriate material for bone replacement scaffolds and has been altered with tyrosine-derived amino acids to make it biodegradable. Polydioxanone has been tested for biomedical applications, while PMMA is commonly used in orthopedic surgery as bone cement for implant fixation and to repair fractures and bone defects such as osteoporotic vertebrae. However, the polymerization of toxic monomers to form PMMA generates significant heat that can harm tissues. Moreover, PMMA is non-degradable and bio-inactive, and it does not chemically bind to bone, resulting in fragmented debris that may cause inflammatory foreign body reactions [92,100]. Other non-degradable polymers commonly used in orthopedic procedures include low-density PE, HDPE, and ultra-high molecular weight PE (commonly utilized as the surface of artificial hip implants [101,102]). Additionally, PU, polyethylene terephthalate, and PP are employed for knee ligament repair [103]. Polyactive™, which is a block copolymer of PEG and PBT, has also undergone research for biomedical applications [104,105,106]. Cellulose [107,108] and its derivatives [109,110] are commonly used in biomedical applications. Furthermore, polyethylene oxide, PHB, and their blends have undergone testing for biomedical purposes [29].

Linear aliphatic poly(α-hydroxyesters), like PLA [111], PGA [112], and their copolymers, are the most commonly used synthetic polymers in medical practice. Manipulation of mechanical properties can be achieved through the polymerization of d-lactide, l-lactide, d,l-lactide, or meso-lactide due to the chiral nature of lactide monomers. Additionally, PLA/PGA copolymers, known as PLGAs, can be tailored to exhibit specific properties, like crystallinity and degradation rate, by altering their composition. These thermoplastic polyesters are fully biodegradable and derived from renewable sources like starch, corn, and sugar cane. They currently rank second in global bioplastic consumption, finding extensive use in both general and specialized applications. Notably, these synthetic polymers have received FDA approval and are the only known biodegradable ones with this distinction [29,93,111,112]. They are biocompatible, primarily non-inflammatory, and degraded in vivo through hydrolysis and enzymatic action, resulting in the excretion of products from the body via normal metabolic pathways [88,93,113]. Additionally, they have potential for drug delivery [114]. Poly(α-hydroxyesters) have been studied as medication and protein delivery systems, including for growth factors, and as scaffolds for tissue regeneration and replacement. They have also been studied as cell carriers. Additionally, they have been used in the production of membranes and films, as well as screws, pins, and plates for orthopedic applications [88,93,115,116]. Among the resorbable polyesters, PGA exhibits a more rapid degradation rate compared to others (less than 12 months) [112]. Therefore, the degradation rate of PLGA can be modified by adjusting the amounts of two-component monomers. In orthopedic applications, this can enable the creation of materials that degrade in sync with bone growth [117]. The varying degradation timeframes present a range of possibilities for biomedical applications. Furthermore, PLGA has been shown to stimulate osteoblast migration and proliferation, which is vital for bone tissue regeneration [93,118]. However, PLGA alone may decrease the impact of stress shielding and is too weak to support a load-bearing environment, making it suitable for specific clinical indications such as ankle and elbow fractures [113]. Additionally, it facilitates self-degradation, resulting in decreased mechanical properties and lowered solution pH, which subsequently leads to extensive degradation. The extensive amount of implant degradation products overwhelm the body and can lead to inflammatory foreign reactions. Ultimately, poly(α-hydroxyesters) lack both bioactivity and osteoconductivity [93,119].

Biomedically relevant polymers can be categorized into synthetic and biogenic types. Synthetic polymers, including PE, PMMA, PLA, PGA, and PCL, are one type. Conversely, biogenic polymers are made up of different kinds of polysaccharides, including cellulose, starch, alginate, chitin/chitosan [120,121], pectin, gellan gum, hyaluronic acid, and their derivatives. Proteins like collagen, fibrin, soy, gelatin, silk, and various bio-fibers [122,123], such as lignocellulosic natural fibers, fall under this category. It is worth noting that natural polymers tend to have a highly organized structure. Extracellular substances named ligands are also possible components of polymers, with the purpose of cell receptor binding. However, natural polymers come with various impurities that must be eliminated prior to use. Synthetic polymers, on the other hand, can be created under controlled circumstances, possessing predictable and reproducible mechanical and physical traits, such as tensile strength, modulus of elasticity, and degradation rate. Synthetic polymers have the potential for chemical and molecular modification, allowing for easy customization to meet specific application requirements while also providing control over impurities. Certain researchers distinguish between non-resorbable polymers, like PE, PP, PMMA, and cellulose, and resorbable or biodegradable polymers, like proteins, poly(α-hydroxyester), and polysaccharides [123]. Synthetic polymers can be categorized as thermoplastics or thermosets. Examples of thermoplastics include HDPE and PEEK, while polydimethylsiloxane and PMMA are examples of thermosets [82]. Table 1 in reference [123] provides a list of synthetic biodegradable polymers utilized as scaffolding materials in biomedical applications. Additional information on polymers suitable for biomedical applications can be found in other sources [82,116,124,125,126,127,128,129]. Excellent literature reviews on the synthesis of diverse biodegradable polymers [130] and current trends in polymer composites [131] are available.

3.3. Inorganic Materials, Substances and Compounds

3.3.1. Metals

Metals are typically assessed for their mechanical performance in bone tissue engineering, despite the potential for increased stress shielding. Titanium and its alloys are among the most biocompatible metals and are widely utilized for bone graft fabrication. Additionally, a range of other metals and their alloys, including zirconium, hafnium, vanadium, niobium, tantalum, chromium, iron, cobalt, nickel, copper, silver, magnesium, zinc, and stainless steel, are employed in the human body [132,133,134,135,136]. Recent research has shown that porous metals have significant biomedical potential [137,138,139]. Metal implants exhibit the necessary strength and toughness to support load-bearing body parts and will continue to play a crucial role as orthopedic biomaterials despite the potential risks of poor wear resistance, as well as the potential release of certain ions and corrosion products from metal implants. Elemental metals do not exist in the human body, thus, neither metals nor their alloys can be considered biomimetic in terms of chemical composition. The term ‘biomimetic’ refers to a processing technique that imitates or is inspired by biological mechanisms [140]. Though biocompatible metals are tolerated by the body and not rejected, they do not actively interact with surrounding tissues. Metal implants frequently exhibit poor attachment to host tissue, thereby limiting their integration with bone. However, in certain scenarios (particularly when coated with CaPO₄ [47]), metallic implants can exhibit acceptable biocompatibility [141]. Traditionally, permanent implants were exclusively made of stable, inert metals or alloys, resistant to degradation and corrosion. However, the paradigm has recently shifted to propose biodegradable implants consisting of Mg, Zn, Fe, and their alloys, allowing for bone ingrowth [142,143].

3.3.2. Glasses and Glass–Ceramics

Special glasses [144,145], glass ceramics [146,147], and pure silica (SiO₂) are suitable materials for biomedical applications. It is recommended to focus on the bioactive ones, which have compositions belonging to the CaO–P₂O₅–SiO₂–MgO–Na₂O system and may have small amounts of CaF₂, MgF₂, K₂O, and SrO [148]. The specific Na₂O–CaO–SiO₂–P₂O₅ formulation known as Bioglass^® [149,150] is the most well-known among them. They are manufactured through standard glassmaking techniques and necessitate pure raw materials. Bioglass^® and other bioactive glasses, which are biocompatible and osteoconductive biomaterials, bind to bones through fibrous connective tissue interfaces, making them extensively employed in bone defect fillings owing to their characteristics. The primary drawbacks of bioactive glasses are their mechanical fragility and low fracture durability accounted for by the amorphous 2D glass network. The flexural strength of most bioactive glass formulations falls within the 40–60 MPa range, rendering them unsuitable for high-load applications. Porous Bioglass^® scaffolds, however, exhibit improved bioactivity and resorption properties [149,150].

Through the process of heat treatment, glasses can be transformed into glass–crystal composites. These composites consist of crystalline phases that are precisely controlled in size and composition. The mechanical properties of the resulting glass–ceramics can often be better than those of the parent glass or sintered crystalline ceramics [146,147]. Bioactive A-W glass–ceramics are produced by a conventional melt quenching process and are made from parent glasses of the pseudo-3CaO·P₂O₅–CaO·SiO₂–MgO·CaO·2SiO₂ system. The bioactivity of A-W glass–ceramics surpasses that of sintered HA. Moreover, their exceptional mechanical properties make them suitable for clinical implementation as prosthetics for iliac and vertebral joints, as well as intervertebral spacers [151,152]. These ceramics serve as a valuable alternative to conventional biomaterials.

3.3.3. Ceramics

Metal oxide ceramics are a category of inorganic biomaterials that have played a significant role in bone tissue engineering. There are a number of factors that support their prominence as biomaterials for bone tissue engineering. Alumina (Al₂O₃), zirconia (ZrO₂), and titania (TiO₂) are among them and have been extensively researched due to their exceptional tribological qualities, bioinertness, high wear resistance, fracture toughness and strength, and comparatively low friction [153,154]. The cooling of pure zirconia from tetragonal to monoclinic phase results in a volume change that creates cracks in zirconia ceramics [155]. To stabilize the material in the tetragonal or cubic phase, it is necessary to mix zirconia with additives such as magnesia (MgO), calcia (CaO), and yttria (Y₂O₃). These materials are commonly known as PSZ [156,157]. However, the clinical application of all ceramics is limited by their brittle nature, necessitating further research to enhance their properties.

3.3.4. Carbon

Elemental carbon has served as a biomaterial since 1972 because of its strength, bioinertness, superior tribological qualities, fracture toughness, and low friction [158]. Its uses include bone bridges, prosthetic hip joints, structural skeletal extensions, glassy carbon roots for artificial teeth, and orthopedic prostheses. Researchers have also studied the biomedical properties of amorphous carbon [159]. However, current research trends focus on biomedical applications of nanotubes and other allotropes of nanodimensional carbons [160,161].

Nanodimensional allotropes of carbon, namely nanotubes, fullerenes, and graphene, exhibit promising potential for biomedical applications because of their small dimensions, high aspect ratios (length/diameter or surface/thickness), large surface area, and excellent mechanical properties. These properties include extreme flexibility and strength, high elasticity, great resistance to bending (in the case of nanotubes), and an ability to undo tube buckling (in the case of nanotubes). Research indicates that non-functionalized nanodimensional carbon tends to aggregate and form bundles while having some biological activities [162,163]. Although these forms are insoluble in water and organic solvents, chemical functionalization [164] allows for better dispersion of carbon nanotubes and improves interfacial bonding within the composite [165]. Technical abbreviations will be explained when first used. Consistent citation and footnote styles will be used throughout the paper. Furthermore, it has been found that the surface functionalization of carbon nanotubes with carboxyl groups can cause calcification, much like in woven bone [166]. CDHA can be deposited in situ on the surface of carbon nanotubes to functionalize them [167]. Surface functionalization is also widely used to enhance the characteristics of other types of nanodimensional carbon, such as fullerenes [168] and graphene [169].

4. A Brief Information on Preparation Techniques

In general, the preparation of composite materials involves wetting, mixing, saturating, and compressing the reinforcement and matrix. The matrices are then bonded together (through heat or chemical reaction) to form a rigid structure. Typically, these processes are performed in open or closed molding dies and involve melting. Different materials require different processing conditions, such as temperature and type of solvent, as well as the order and method of adding components vary greatly. Depending on the constituents, composites can be produced in a variety of ways, including fiber placement, filament winding, souring, tufting, z-pinning, various types of casting and molding, braiding (into formers), filament winding, and slip molding [38]. However, before mixing, the particle size dimensions of the mixed components need to be considered. CaPO₄, carbon, and metal oxide ceramics are almost always fine powders that can be mixed immediately, but this is not the common case for other materials. For example, polymers are often supplied as pellets, granules, or coarse powders and often require an additional grinding step, and, for low-melting polymers such as PCL, cryogenic milling can be effective. Similarly, glasses and metals may also require a grinding process.

With regard to CaPO₄-based formulations, various methods have been realized to combine matrix and dispersion components to form biocomposites. For example, mechanical blending or mixing, ball milling, compounding, compression, and casting into a polymer–solvent solution followed by solvent evaporation, e.g., by freeze-drying, dispersion of CaPO₄ particles, or whisking into a liquid monomer, followed by polymerization, a melt extrusion of CaPO₄/polymer mixtures, and co-precipitation or co-deposition, as well as a rapid prototyping, selective laser sintering, and 3D printing [22,37,170,171,172,173]. Four-dimensional printing, in which the resulting 3D shape is able to morph into different forms in response to environmental stimulus, with the fourth dimension being the time-dependent shape change after the printing, is used as well [174]. There has been a comparison of three techniques for creating homogenous blends of PLLA and HA [170]. Prior to compression molding, polymer pellets and ceramic powder were combined using a dry technique. The second technique involved distributing the ceramic filler into a solution of polymer and solvent. The third technique used combined powders of ceramic and polymer via melt extrusion. A network of ceramic particles surrounds the polymer pellets when a dry powder is combined, but the solvent and melt procedures also produce a uniform distribution of HA in the matrix. The possibility of hazardous organic solvent remains is the primary drawback of the solvent casting technique. It has been demonstrated that the melt extrusion approach is appropriate for creating homogenous ceramic/polymer blends [170]. However, it should be noted that non-melt-based routes often lead to the development of composites with poor mechanical performance and often require the use of toxic solvents and intensive manual labor [125].

Other options include in situ formation, where the reinforcement is synthesized within a pre-formed matrix, or synthesizing the matrix around the reinforcement [37,175,176,177,178]. One of the most appealing approaches is this one since it prevents large-scale particle aggregation. For instance, a number of studies have described in situ creation methods for creating different composite materials with carbon nanotubes and CaPO₄ [179,180,181,182]. The same is true for CaPO₄/graphene composites [177,178]. Other examples include biomimetic synthesis [177,178,183,184], the use of amino acid-coated gold nanosized particles as scaffolds to grow CDHA [185], and the preparation of nanosized HA/PA biocomposites [186,187]. In some cases, mechanochemical routes [188,189], emulsions [190,191,192,193,194,195], lyophilization [171,196,197] and freeze–thaw techniques [198], and gel templating mineralization [198,199], as well as vat photopolymerization [200], can be applied to produce CaPO₄-based biocomposites. Various techniques, such as in situ preparation, spark plasma synthesis, hydrothermal treatment, biomimetic mineralization, hot isostatic pressing, electrochemical deposition, and ball milling, have been used to prepare graphene/HA nanocomposites [178]. In the case of CaPO₄ biocomposites with metals, a powder metallurgy approach [201,202] combining direct ink writing with liquid pressure infiltration [203], as well as various types of additive manufacturing [204,205], can be used. The details of various production methods are well described in other publications [22,37,117,170,171]. Furthermore, additional types of preparation techniques are briefly mentioned below when describing the individual CaPO₄-based formulations.

After the desired formulations of CaPO₄-based composites and hybrid formulations have been prepared, it frequently becomes necessary to create 3D constructs from them, which might be either the standard geometric shapes (cubes, rectangular blocks, cylinders, filaments, etc., as well as various types of screws and dowels) or complicated shapes of the specific bones or bone fragments. The detailed description of the forming and shaping techniques is beyond the scope of this review but, briefly, one can mention the following. Depending on the nature of the second phase of the biocomposites (CaPO₄ is considered as the first one), the standard geometric shapes might be created by common ceramic processing techniques [206,207], metal processing techniques [208,209,210], and/or polymer processing ones [123,211], while the complicated shapes are commonly created by various types of additive manufacturing techniques, such as 3D printing [173,212,213,214,215].

Commonly it is necessary to impart porosity to CaPO₄-based composites and hybrid formulations, which is advantageous for most applications as a bone replacement material because porosity facilitates the migration of osteoblasts from the surrounding bone to the implant site. Various material processing strategies to prepare composite scaffolds with interconnected porosity include thermally induced phase separation, solvent casting, particle leaching, solid free-form fabrication techniques, microsphere sintering, and coating [123,216,217,218,219]. Supercritical gas foaming techniques can also be used [170,220,221,222]. Finally, different types of additive manufacturing techniques can also be used [212,213,214,215]. Most of those porosity-forming techniques have been adapted from (bio)ceramics [223,224]. Regarding CaPO₄-based biocomposites with polymers, the most common technique to produce porous scaffolds appears to be fused deposition modeling (64%), followed by low-temperature deposition manufacturing (19%), selective laser sintering (12%), and stereolithography (5%) [225]. Further details on porosity creation are available in the topical reviews [218,223].

5. CaPO₄-Based Biocomposites and Hybrid Biomaterials

There are multiple (partially overlapping) major categories into which CaPO₄-containing composites and hybrid formulations appropriate for biological applications can be classified:

• Biocomposites with polymers;
• Self-hardening biocomposites;
• Nanosized CaPO₄-based formulations and nanosized biocomposites;
• Collagen-containing biocomposites;
• Biocomposites with other bio-organic compounds;
• Injectable bone substitutes (IBS);
• Biocomposites with glasses, inorganic compounds, carbon, and metals;
• Biocomposites from CaPO₄ only;
• Inks for 3D printing;
• Functionally graded biocomposites;
• Biosensors.

The majority of these have been developed using bone-like concepts to mimic natural bone. Details on each of these areas are as follows.

5.1. Biocomposites with Polymers

5.1.1. History and General Part

Since natural bones represent a biocomposite of ion-substituted CDHA with a bioorganic polymer collagen [18,19], it is plausible to believe that an appropriate biocomposite with mechanical, chemical, and physical properties similar to those of human bones would be made of CaPO₄ (such as CDHA) and a bioorganic polymer (such as collagen). Therefore, CaPO₄/polymer composites and hybrid formulations have been widely investigated to improve the properties of the constituents with the final purpose of bone regeneration applications. In such biocomposites, hard and rigid CaPO₄ bioceramics provide the basic building blocks for both mechanical strength and biomineralization, while flexible and soft polymeric components provide a necessary level of elasticity. Depending on the amounts of the constituents, such formulations can be broadly classified into two categories: CaPO₄-reinforced polymers and polymer-reinforced CaPO₄ bioceramics. Therefore, the formation of CaPO₄/polymer composites and hybrid formulations takes advantage of the advantages of both materials and minimizes their disadvantages. In general, the elastic modulus of polymers is lower (up to 2–7 GPa) than that of bone (3–30 GPa), which means that CaPO₄ bioceramics must be loaded at high weight ratios to mimic the latter. General knowledge of composite mechanics also shows that particles with high aspect ratios, such as whiskers and fibers, significantly increase the modulus at low loads. Therefore, some attempts have been made to prepare biocomposites containing whisker- [226,227,228,229,230] or needle-shaped [231,232,233,234] CaPO₄ and CaPO₄ fibers [235].

Implantable CaPO₄/polymer biocomposites have a history that dates back to 1981 (while the subject of “ceramic-plastic materials as bone substitutes” is at least 18 years older [236]), with pioneering research on HA/PE blends initiated by Professor William Bonfield and colleagues at the Queen Mary and Westfield College, University of London, UK [237,238]. That early work introduced the concept of bone resemblance, where the proposed biocomposite was composed of a polymeric ductile matrix of PE and a ceramic rigid phase of HA. Of them, HA stiffened the polyethylene, while PE toughened the composite. That approach was significantly extended and developed in further research by the same group [66,239,240,241,242,243,244,245]. Further studies investigated the effect of the surface topography of HA/PE composites on cell proliferation and attachment [246,247,248,249]. The material consisted of a specific combination of volume-loaded 40% HA particles homogeneously dispersed in an HDPE matrix. The idea was to mimic bone by reinforcing the polymer matrix, which can develop significant anisotropy with appropriate orientation techniques, with bone-like bioceramics that guarantee both mechanical reinforcement and bioactive properties of the composite material. Following FDA approval in 1994, a HA/PE formulation with 40 vol.% HA was given a trade name HAPEX™, became commercially available in 1995 [243,244,245,247] (Smith and Nephew Richards, Bartlett, TN, USA), and has been successfully implanted in more than 300,000 patients to date. It represents a major step forward in the bone grafting field because it was the first commercial product of the ‘second-generation’ also called surface-active biomaterials [250] that have been developed to be bioactive rather than bioinert [251]. The three primary steps in the production of HAPEX™ include blending, compounding, and centrifugal milling. Bulk materials and devices are then made from this powder through compression and injection molding [37]. In addition, HA/HDPE biocomposites can be prepared by hot rolling techniques that facilitate homogeneous dispersion and mixing of the reinforcement within the matrix [252]. Additionally, PP can be used instead of PE [253,254,255,256].

Mechanical interlocking between the two phases of HAPEX™ is formed by the shrinkage of HDPE on HA particles during cooling [66,67,257] and both HA particle size and its distribution in the HDPE matrix have been identified as important parameters affecting the mechanical behavior of HAPEX™. Namely, smaller HA particles were found to result in stiffer composites due to an overall increase in the interface between the polymer and the ceramic. Furthermore, the stiffness of HAPEX™ was found to be proportional to the HA volume fraction [242]. The use of coupling agents, such as 3-trimethoxysilylpropyl methacrylate, for HA and acrylic acid for HDPE can improve the bonding (through both chemical and mechanical adhesion) between HA and HDPE [258,259]. Naturally, in biocomposites including PE, alternative CaPO₄ compounds can be employed in place of HA [260]. Additionally, in an effort to enhance the mechanical qualities, efforts have been made to incorporate other bioceramic phases into the HDPE matrix of HAPEX^TM, such as PSZ [261] and alumina [262]. For example, replacing a portion of the HA filler particles with PSZ ones was found to increase the strength and fracture toughness of HA/HDPE biocomposites: The compressive stresses generated by the volume expansion associated with the transformation of PSZ from tetragonal to monoclinic phase inhibit or delay crack propagation in the composite. As a result, the fracture toughness of HA/ZrO₂/HDPE biocomposites is increased [261].

Numerous investigations have demonstrated that HAPEX^TM adheres to bone directly via chemical bonding (biological fixation) as opposed to fibrous encapsulation (morphological fixation); orbital reconstruction was the initial clinical application of HAPEX^TM [263], but, since 1995, the biocomposite’s primary use has been in the sound transmission shafts of middle ear implants (Figure 1 [264]) for the purpose of treating hearing loss [265,266]. The advantage of HAPEX^TM in both applications is its ability to be molded in situ, which gives surgeons the opportunity to make final adjustments to optimize the prosthesis’ fit to the patient’s bone. This means that there is minimal risk of failure due to insufficient tensile strength for subsequent activities, and only limited mechanical loading is needed [66,67,149]. When HA concentrations are less than 40%, HA/PE composites have superior fracture toughness compared to cortical bone, while concentrations between 45 and 50% show comparable fracture toughness. The Young’s modulus result is in the range of 1–8 GPa, which is very close to that of bone. Observation of the fracture surface shows that only a mechanical bond is formed between HA and PE. Unfortunately, HA/PE composites are not biodegradable, the available surface area of HA is low, and the presence of bio-inert PE reduces its ability to bond to bones. Furthermore, HAPEX™ is designed at maximum density to increase strength, which results in a lack of void space limiting osteoblast viability when the implant is placed in the body [23,150]. More information about HAPEX™ can be found in the literature [66,67]. In addition, other types of HA/PE biocomposites have been developed [267,268,269,270,271,272,273].

PE has been utilized in both linear and branching forms as the matrix; biocomposites made with the former have higher moduli [270]. There is currently no compelling explanation for the mechanisms of reinforcement in CaPO₄/polymer compositions. Generally speaking, if the filler is selected poorly, the polymer matrix will be impacted by it. This can be seen in the forms of inert polymer shell formation around the particles (surface-induced crystallization, transcrystallization, and epitaxial growth), a decrease in molecular weight during composite processing, and conformational changes in the polymer caused by the interparticle space and particle surface [66,67]. On the other hand, the molding method may have an impact on the reinforcing effect of the CaPO₄ particles. It has been discovered that the composite’s mechanical performance increases with the orientation of the polymer matrix [267,268].

It is also possible to create a wide range of alternative CaPO₄ mixes with other polymers [274], including very uncommon formulations that comprise dendrimers [275]. Moreover, CaPO₄/polymer compositions with light curing are known [276]. Table 3 provides a list of acceptable CaPO₄ products (MCPM and MCPA are both highly acidic and hence not biocompatible [16]; however, this drawback can be addressed by combining them with basic compounds such as TTCP, HA, CaO, and CaCO₃). There are two purposes of combining CaPO₄ with polymers to make biocomposites: the desirable mechanical properties of the polymer compensate for the poor mechanical behavior of CaPO₄ bioceramics and, conversely, the desirable bioactive properties of CaPO₄ enhance those of the polymer, expanding the possible uses of each material in the body [277,278,279,280]. Namely, the addition of polymers to CaPO₄ increased its mechanical strength [277] and the addition of CaPO₄ filler to polymers increased their compressive strength and modulus and osteoconductive properties [119,281,282,283,284]. As CaPO₄ content increased in biocomposites, Young’s modulus and bioactivity generally increased, and ductility decreased [23]. Subsequent research has revealed that the mechanical characteristics of CaPO₄/polymer biocomposites are more complex. It was discovered that the strength of these biocomposites decreased as the CaPO₄ level increased [285]. However, the biocompatibility of such biocomposites increases as the CaPO₄ filler causes an increased initial flash spread of serum proteins compared to more hydrophobic polymer surfaces [286]. Furthermore, experimental results of those biocomposites showed enhanced cellular activity and good cell–material interactions compared to the corresponding polymers alone [279]. Moreover, these compositions have the ability to release calcium and orthophosphate ions into myeloma in a sustainable manner, both of which are critical for the regeneration of mineralized tissue [278]. In fact, the combination of two different materials can produce better biocomposites than the materials alone, taking advantage of the strengths of each. However, in some cases, the incorporation of CaPO₄ into biocompatible, mechanically stable, and 3D printable polymers does not improve bone formation in vitro or in vivo [287].

5.1.2. Apatite-Based Formulations

It is recognized that the primary inorganic phase of mammalian calcifying tissues is biological apatite [14,15]. Therefore, CDHA, HA, carbonate apatite, and sometimes FA (all of which can be both pure and contain dopants) have been applied in the preparation of biocomposites with other polymers, often with the aim of improving their biological activity. For example, polysulfone composed of HA can be used as a starting material for long-term implants [288,289,290]. In in vivo experiments, HA/polysulfone biocomposite-coated specimens from the distal femur of a rabbit showed direct bone attachment to the coating compared to the fibrous encapsulation that occurs when using uncoated specimens [288]. The bioresorption time of such biocomposites is a very important factor that depends on the microstructure of the polymer and the presence of modified phases [289].

Various apatite-containing biocomposites have been developed using PVA [198,291,292,293,294], PVAP [295], and other polymers [296,297,298,299,300,301,302,303,304,305,306]. Namely, PVA/CDHA biocomposite blocks were prepared by precipitating CDHA in an aqueous solution of PVA [198]. An artificial cornea consisting of a porous skirt of nanoscale HA/PVA hydrogel and a transparent center of PVA hydrogel was also created. The results showed good biocompatibility and docking between the artificial cornea and the host tissue [291,292]. PVAP was chosen as the polymer matrix due to its phosphate groups acting as binding/anchoring agents and its high affinity for the HA surface [295]. HA/Ca poly(vinyl phosphonate) biocomposites were also created [299,300], while a high-affinity integration of CDHA with PHEMA hydrogel scaffolds was studied by template-guided nucleation and mineral growth processes [305].

Biocomposites including HA have also been created using PEEK [226,227,307,308,309,310,311,312] and a high-impact polystyrene [313,314], which may find therapeutic application in load-bearing applications. Mechanical properties increased monotonically with reinforcement concentration in tests where PEEK was reinforced with thermally sprayed HA particles; the highest values were obtained in studies where the volume fraction of HA particles was less than 40% [307,308]. The stated strength range of 45.5–69 MPa and hardness range of 2.8–16.0 GPa surpass the lower values for human bone, which are 7–30 GPa and 50–150 MPa, respectively [308]. Since the tensile strength of conventional PEEK/HA biocomposites fell sharply with increasing HA loading, nanodimensional HA rods and carbon nanofibers were used to reinforce PEEK. In that case, the tensile properties of a PEEK/15 vol% HA/1.9 vol% carbon nanofibers biocomposite were found to compare favorably with those of human cortical bones [310]. Modeling of the mechanical behavior of HA/PEEK biocomposites is available elsewhere [309].

Biodegradable poly(α-hydroxyesters) are well known in clinical medicine. Currently, these materials offer good options when looking for suitable polymer fillers. For example, cytocompatible HA/PLGA formulations have been developed for bone tissue regeneration [315,316,317,318,319], while highly porous PLLA foam was seeded with HA particles to enhance the osteoconductivity of polymer scaffolds for bone tissue engineering [281]. Since hydration increases the quantity of COOH and OH groups on the polymer surface, which serve as apatite nucleation sites, the foam’s amount of carbonized CDHA material rose before it was incubated in simulated bodily fluid. More information about the mechanical characteristics of PLA/CaPO₄ biocomposites made using various methods, as well as the outcomes of in vitro and in vivo tests, is available elsewhere [316].

Poly(α-hydroxyesters), such as PGA and PLA, are known to spontaneously decompose into acidic monomers (glycolic acid and lactic acid, respectively), which catalyze further degradation and cause inflammatory reactions in surrounding tissues [320]. Therefore, in CaPO₄ biocomposites with poly(α-hydroxyesters), the presence of slightly basic compounds (HA, TTCP) neutralizes the acid molecules to some extent, leading to a weak pH buffering effect on the polymer surface, thus partially compensating for these disadvantages [119,317,321,322,323]. However, the addition of more basic chemicals (e.g., CaO, CaCO₃) may become necessary [123,323,324,325]. Extensive cell culture experiments on pH-stabilized PGA and carbonate apatite composites have been reported and subsequently supported by extensive in vitro pH studies [326], which resulted in the design of functionally graded composite cranial implants composed of PLA, carbonate apatite, and CaCO₃ [327]. Except for the pH buffering effect, the addition of CaPO₄ also increases the hydrophilicity and water absorption of the polymer matrix, thereby modifying the degradation kinetics of the scaffold. For instance, when water enters the interface region, it has been observed that polymer biocomposites containing HA particles hydrolyze uniformly [328].

Poly(α-hydroxyester) and CaPO₄ biocomposites are mainly prepared by incorporating the inorganic phase into a polymer solution followed by drying under a vacuum. The resulting solid biocomposites can be molded using various processing techniques. These biocomposites can be made by combining l-lactide and HA particles prior to polymerization [321] or by combining slip casting and hot pressing [329], while other methods of production are also known [316,318,330,331]. The addition of surface active agents (surfactants) may be useful to maintain the homogeneity of the suspension [332]. It is also possible to create HA/PLA [191,192] and HA/PLGA [193] microspheres using microemulsion technology. There are other known more intricate formulations, such as carbonated-FA/PLA [333] and PLGA/carbon nanotube/HA [334]. Durucan and Brown have released a collection of interesting references pertaining to various ways of manufacturing HA/poly(α-hydroxyester) biodegradable composites [335,336,337]. The authors prepared CDHA/PLA and CDHA/PLGA biocomposites by solvent casting method, followed by hydrolysis of α-TCP to CDHA in aqueous solution. It was found that the presence of both polymers inhibited the hydrolysis of α-TCP compared to single-phase α-TCP and the inhibitory effect of PLA exceeded that of PLGA [335,336,337]. The physical interactions between poly(α-hydroxyester) and CaPO₄ are shown in Figure 2 [337]. Another good illustration is available in the literature [51].

CaPO₄ addition has the potential to greatly improve the mechanical characteristics of poly(α-hydroxyesters) [338,339]. Therefore, fixations (screws and plates) built from those composites have been created and evaluated, and CDHA/PLLA biocomposites with extremely high mechanical properties have been developed [119]. The fixations have demonstrated easy handling and molding to the shape of the implant site, complete absorbability, excellent ability to bond directly to bone tissue without fiber tissue, osteoconductivity, biocompatibility, and high stiffness that can be maintained for the time required for bone union [328,330]. Elastic modulus can reach 12 GPa, and the initial flexural strength of 280 MPa is more than that of cortical bone (120–210 MPa) [119]. In phosphate-buffered saline, this strength could be sustained above 200 MPa for a maximum of 25 weeks. These biocomposites were made by precipitating PLLA/dichloromethane solutions that contained scattered tiny CaPO₄ particles [118,340]. PDLLA + HA [221,341,342] and PLGA + HA [343], along with the multicomponent compositions PLA/ethylcellulose/HA [344] and PLA/Fe₃O₄/BCP (HA + β-TCP) [345], were used to fabricate biocompatible porous scaffolds. Compared to single-phase HA bioceramics, biodegradation was markedly accelerated, and freshly generated bone was seen when implanted in rabbit femurs. This could be because HA dissolves when lactic acid is released locally. PLA and PGA fibers were attached to porous HA scaffolds in additional research. The fact that the reinforcement did not prevent bone from penetrating the implant encourages the development of these biocomposites further for usage as alternatives to bone grafts [29,316,317]. Recent data on using of 3D-printed PLA/CaPO₄ composite scaffolds for bone tissue engineering in preclinical in vivo studies are available in the literature [346].

Biocomposites, termed SEVA-C, were produced by combining EVOH-starch blends with 10–30 wt% HA fillers. The samples that contained 30 wt% HA had moduli up to 7 GPa, as reported in [347,348,349,350,351]. The purpose of adding bioactive HA fillers in SEVA-C was to create composites that exhibit bioactivity and possess the necessary stiffness similar to those found in human cortical bone. Results indicated that these biocomposites exhibited significant bioactivity in vitro, which was supported by the polymer’s water uptake capacity [352]. However, the addition of HA particles to SEVA-C was determined to impact the rheological properties of the blends. A degradation model was subsequently constructed for these mixtures [353].

The homologs poly(3-hydroxybutyrate) and poly(3-hydroxyvalerate) are not very biodegradable. However, biocomposites of these polymers with CaPO₄ have shown good biocompatibility both in vitro and in vivo [354,355,356,357,358]. The bioactivity and mechanical properties of these biocomposites can be tuned by changing the volume fraction of CaPO₄. Similarly, biocomposites of PHBHV with HA and amorphous carbonate apatite (almost ACP) have been shown to have promising potential for repair and replacement of damaged bone [359,360,361,362].

In this framework, a less biodegradable but biocompatible polymer, PCL, has been employed, and PCL/HA and PCL/CDHA biocomposites have been considered viable options for bone tissue regeneration and repair [216,363,364,365,366,367,368,369]. For example, biocomposites were obtained by infiltrating porous apatite blocks with ε-caprolactone monomers and polymerizing them in situ [364]. The composite was found to be biodegradable and potentially applicable as a bone replacement and cartilage regeneration material for cancellous bone and trabeculae; the addition of HA strongly enhanced both the mechanical performance and biocompatibility of PCL in osteoblast cultures [370]. For the preparation of PCL/HA biocomposites, various techniques are known [216,366]. For example, to make PCL and nanosized HA biocomposite fibers, the desired amount of nanosized HA powder was dispersed in a solvent using a magnetic stirrer followed by sonication for 30 min. PCL was then dissolved in this suspension and the solvent was evaporated [371]. The reverse preparation sequence is also possible: PCL was dissolved in chloroform at room temperature (7–10 wt%/volume), HA (~10 µm particle size) was suspended in this solution, sonicated for 60 s, and then the solvent was evaporated [372] or immersed in salt [373]. The mechanical properties obtained with this technique are approximately one-third of those of cancellous bone. In a comparative study, PCL and bio-apatite were mixed in an extruder in a 19:1 ratio [374]. After preparation was complete, the mixture was cooled in a nitrogen atmosphere. The authors observed that the presence of bio-apatite increased the modulus of elasticity and also increased the hydrophilicity of the polymeric substrate. Furthermore, increasing the apatite concentration was found to improve both the elastic modulus and yield stress, indicating a good interfacial interaction between biological apatite and PCL. The presence of biological apatite was also found to stimulate osteoblast attachment to the biomaterial and cell proliferation [374]. In another study, PCL/HA biocomposites were prepared by melt stirring at 120 °C in a rheometer connected to a stirrer until the torque reached equilibrium [375]. The samples were then compression molded and cut into specimens of suitable size for testing. The results showed that composites containing 20 wt% HA had the highest strength [375]. However, grafting PCL directly onto the surface of HA particles seems to be the most interesting preparation method [363]. In one more study, HA porous scaffolds were coated with a PCL/HA composite coating [376]. In that system, PCL as a coating component improved the brittleness and low strength of the HA scaffold, while the particles in the coating improved the osteoconductivity and bioactivity of the coating layer. In addition, there are multicomponent biocomposites, including PDLLA/PCL/HA [377], PLLA/PCL/HA [378], FA-HA/PCL [379], chitosan/PCL/PVA/HA [380], magnetic PCL/Fe-doped HA [381], supramolecular PCL/functionalized HA [382,383], Cu₂O@CuFeS₂/HA/chitosan [384], etc. In November 2023, results of the first clinical case of the Ilizarov method of bone reconstruction by means of a bioactive and degradable intramedullary HA/PCL biocomposite implant were published [385]. More details on PCL/HA biocomposites and processing methods can be found elsewhere [216,366,368].

At the end of this section, HA biocomposites with other polymers need to be mentioned [386,387,388,389,390,391,392,393,394,395,396]. Among them, PTMC/HA biocomposites have been suggested to be suitable for stereolithographic fabrication of patient-specific flexible implants [393]. Moreover, some of them exhibit interesting properties. Namely, polymers with shape memory properties are worth mentioning [397]. Consequently, it is also thought that composites based on CaPO₄ and these polymers exhibit shape memory qualities [389,390,391]. More complex formulations are also known. Namely, biopolymer/CHDA biocomposites encapsulating magnetic nanoparticles [398], as well as many other magnetic formulations [399], have been prepared. More details on HA/polymer composites and hybrid formulations can be found in other reviews [400,401,402,403].

To conclude, currently (May 2024), according to Scopus, there are 900 articles with a combination of “polymer” and “apatite” in the title.

5.1.3. TCP-Based Formulations

Compared to HA, both α-TCP and β-TCP are more soluble in water (Table 3). This is the reason for their faster bioresorption in vivo (although there have been reports of not biodegrading TCP after implantation in calvarial bone defects [404]). Therefore, both types of TCP are widely used to create fully biodegradable biocomposites [405,406,407,408,409,410,411,412,413,414,415,416,417,418,419,420]. For instance, β-TCP particles and gelatin have been combined to create a biodegradable and osteoconductive biocomposite [409]. When this substance was evaluated in vivo, positive outcomes were seen. The material was found to be biocompatible, osteoconductive, and biodegradable, and did not require another surgical procedure to remove the device after healing. To that composition, one may add K₂HPO₄ [411] and botanical extracts [410]. After creating a biocomposite using cross-linked gelatin and β-TCP, another research team implanted it subcutaneously in rats and saw good biocompatibility as well as bone development [412]. Later, this was extended to a porous (porosity ~75%) β-TCP/gelatin composite material containing BMP-4 [415]. Furthermore, porous β-TCP/alginate/gelatin scaffolds with cytocompatibility and osteoinductive properties were fabricated and successfully tested in vitro [413]. In addition, multicomponent formulations, such as a phytoconstituents-filled β-TCP biocomposite with gelatin and pectin polymers incorporating Cissus quadrangularis L. extract [421], as well as another biocomposite of microporous β-TCP filled with alginate-gelatin crosslinked hydrogel, clindamycin, and bone morphogenetic protein (BMP-2) [422] were prepared. Both formulations appeared to be suitable for drug delivery. In such formulations, CaPO₄ filler was found to have a reinforcing effect [423]. In other studies, β-TCP biocomposites with PLLA [340,405,406,407] and PLGC [408] were prepared. β-TCP was able to resist the acidic degradation of polyesters to some extent, but could not prevent the pH from dropping to ~6. However, the biocomposite was successfully implanted into the mandible of a beagle [408]. Similarly, α-TCP/gelatin biocomposites are also known [409].

Based on the concept of self-reinforcement, TCP and polylactic acid biocomposites were prepared and studied using conventional mechanical tests [424]. Resorbable scaffolds were prepared from those biocomposites [425]. Chitosan was also used as a matrix to incorporate β-TCP through solid–liquid phase separation of the polymer solution followed by sublimation of the solvent. Due to the complexation of the functional groups of chitosan with the calcium ions of β-TCP, those biocomposites had high compressive modulus and strength [426]. PCL/β-TCP biocomposites were also developed in other studies [368,369,427,428,429,430,431], and an immersion in simulated body fluids at 37 °C allowed for systematic monitoring of their in vitro degradation behavior [429]. Loading of drugs into PCL/β-TCP biocomposites is a further development of this subject [368,430].

In vitro studies using primary rat cervical osteoblasts have shown increased cellular activity in BMP-loaded β-TCP/gelatin biocomposites [415]. Other researchers investigated BMP-2 loaded porous formulations (porosity ~95%, average pore size 180–200 µm) [432] and confirmed the results. Good results were obtained after 12 months from long-term implantation research with PDLLA/α-TCP biocomposites in a sheep implant loading model; however, after 24 months, a severe osteolytic response was detected. This resulted from the residual PDLLA and the nearly total disintegration of α-TCP by this point [433].

There exist more intricate CaPO₄-based biocomposites. As an illustration, certain formulations comprise TCP, CDHA, and PLGA, three interpenetrating networks [434]. Firstly, a hydrolyzable α-TCP slurry was applied to PU foam to create a porous TCP network. Second, a self-hardening CaPO₄ formulation inserted into the porous TCP network yielded the CDHA network. Lastly, the remaining open porous network of the CDHA/α-TCP complex was compromised by PLGA infiltration. There were three phases to the biocomposite, and each phase had a distinct degrading tendency. It is thought that bone forms on the PLGA network, which deteriorates the quickest, while the remaining CaPO₄ phase holds its structure and capacity to support weight [434]. Furthermore, PCL/TCP/boron nitride biocomposites were prepared in a similar manner [435].

5.1.4. Biocomposites with Remaining CaPO₄

The number of publications on composites and hybrid formulations containing the remaining CaPO₄ compounds and polymers is considerably smaller than the ones with apatite and TCP. Biphasic calcium phosphate (BCP), a solid blend of HA with β-TCP (similar formulations of HA + α-TCP and α-TCP + β-TCP are also known [436]) seems to be the most popular among the remaining CaPO₄ compounds. Namely, hydrophilic PEG/vancomycin complexes were coated on porous PDLLA/BCP scaffolds to modify their surface and distribute drugs [437]. More specifically, both PLGA/BCP [438,439] and PLLA/BCP [440] biocomposites have been produced and found acceptable for repairing, filling, and growing native bone tissue due to their cytotoxic and fibroblastic properties [441,442]. In addition, PCL/BCP [443,444], PTMC/BCP [445], PMMA/BCP [446], gelatin/BCP [447,448], and gelatin/PVA/BCP [449] biocomposites are known as well.

DCPD-based biocomposites composed of DCPD, albumin, and duplex DNA were prepared by water/oil/water interfacial reaction method [190]. Wet spinning was used in a different work to create core-shell DCPD/chitosan biocomposite fibers [450]. P atoms were found in trace numbers inside the composite fibers, whereas Ca and P atoms were mostly found in the outer layer, according to energy-dispersive X-ray spectroscopy. This showed that the chitosan core and CaPO₄ shell of the composite fibers created a distinctive core-shell structure [450]. For use in self-hardening biocomposites, similar formulations have been produced [451]. Moreover, BSA and DCPA nanoparticles were co-precipitated in ethanol to create DCPA/BSA biocomposites [452], DCPA/ethylcellulose ones were prepared using the sol-gel method [453], while those of DCPA/PVA were produced by freeze-drying [454]. Nanosized particles of DCPA were synthesized and incorporated into dental resin to form dental biocomposites [455,456,457,458], while DCPA/dextran/CMC nanocomposite porous scaffolds were prepared by freeze-drying in a lyophilizer [459]. Interestingly, certain DCPD/polymer composites have the potential to be employed in battery devices as proton conductors [460,461]. Although biocompatibility has not been reported, perhaps in the future better formulations will be used to provide biocompatible batteries for electronic devices that are implanted.

To address the effects of dental caries, a number of hybrid and ACP-based biocomposite formulations have been developed for dental applications [462,463]. Among them, there are rechargeable ones with sustained ion release and re-release, which were introduced in 2016 [464,465,466,467]. Furthermore, various ACP-based formulations have been investigated as potential biocomposites for bone grafting [362,468,469,470] and drug delivery [471,472]. Namely, ACP/PPF biocomposites have been prepared by in situ deposition [469], and PHB/carbonated ACP and PHBHV/carbonated ACP biocomposites appear to be suitable as slow biodegradable bone replacement materials [362].

OCP-based biocomposites with polymers are also known. For example, OCP granules were combined with gelatin and the prepared composites were spread on rat calvaria. Those composites were found to promote bone growth due to the balanced cellular activity of osteoblasts and osteoclasts [473]. There is also an excellent review on OCP/polymer biocomposites with a controlled hierarchical structure from the macro to the nanoscale [474]. The following topical overview contains more details about OCP-based formulations that are relevant to biomedicine [475].

Finally, just a few publications on TTCP-based biocomposites with polymers have been published [476,477].

To conclude, currently (May 2024), according to Scopus, there are 404 articles with a combination of “polymer” and “calcium phosphate” in the title.

5.2. Self-Hardening Biocomposites

The early 1980s saw the introduction of inorganic self-hardening CaPO₄ formulations (cements) that were set in vivo [478]. They have been the subject of in-depth research and numerous formulations since then. These mixtures solidify by a variety of chemical reactions among CaPO₄ and eventually form a monolithic body composed of CDHA or DCPD, with other phases possible to mix. Unfortunately, self-hardening CaPO₄ formulations with ceramic properties become brittle after setting, while the hardening time may not be suitable for clinical procedures [478]. Therefore, several attempts have been made to transform them into biocomposites to improve their properties, for example, by adding PEG [479,480], polylactide [481], PLGA [482], gelatin [414,483,484,485,486], osteocalcin/collagen [487], alginic acid [488], chitin [489], chitosan [451,490], fibrin glue [491], silk fibroin [492,493], silanized HPMC [494], CMC [495], PVA fibers with CMC [496] and without CMC [497], CMC/gelatin/fullerenol [498], bioactive glass [499,500,501,502], bioactive glass functionalized with hypoxia conditioned medium [503], calcium silicate [504,505,506], cockle shell powders [507], metals [508], magnetic nanoparticles [509], and so on. Light-curing formulations have also been prepared [510,511]. Furthermore, a wide range of reinforcing additives with varying forms and characteristics are frequently employed to enhance the mechanical characteristics of CaPO₄ compositions that self-harden [512,513]. For this, carbon nanotubes have even been employed [514,515]. Furthermore, the addition of a carbonate apatite type B powder into a self-hardening α-TCP formulation was also found to improve the mechanical properties of the final product [516]. An important part of the reinforced formulations can be named CaPO₄-based concretes [517,518], although this term is rarely used in the field of biomaterials [478]. The idea with concrete is simple: the presence of a strong filler in the matrix can stop the proliferation of cracks. Most of the aforementioned compounds were added to the self-hardening CaPO₄ formulations with the main purpose of improving either the rheological properties of the liquid formulations or the mechanical properties of the hardened ones; however, other aims are also known. For example, the addition of magnetic nanoparticles was found to increase cell adhesion [509].

Various apatite-containing formulations have been developed using PMMA [446,519,520,521,522,523,524,525] and polyethylmethacrylate [526,527]. Such biocomposites are prepared by dispersing apatite powder in a polymeric viscous liquid [528] and can be used for drug delivery purposes [529]. The mechanical properties of concretes composed of PMMA matrix and HA particles of various sizes were tested and tensile results showed that the strength was independent of particle size. The tensile strength did not change after immersion in Ringer’s solution, but the fatigue properties were significantly reduced. In vivo, tests for PMMA/HA blends’ biocompatibility have validated the implants’ enhanced osteogenic qualities over single-phase PMMA [520,521,522,523]. Not only were the properties improved, but the addition of HA also enhanced the osteoblastic response of PMMA [521]. Thus, the addition of CaPO₄ made the non-biodegradable PMMA more bioactive and osteoconductive, resulting in biocomposites with good workability. On the other hand, PMMA/HA biocomposites have low flexural, compressive, and tensile strengths. More complicated formulations, such as PMMA-based HA/ZnFe₂O₄/ZnO biocomposites with antibacterial performance and low toxicity, were prepared as well [530].

It is believed that biocomposites composed of CaPO₄ and different kinds of resins have mechanical and biological characteristics similar to those of ordinary PMMA cement and can be used for implant fixation [228,229,531,532,533]. To enhance the mechanical characteristics of self-hardening CaPO₄-based formulations and maintain stability at the implant site, numerous researchers have turned to in situ curing formulations, mostly via polymer matrix crosslinking processes [511,534,535,536,537]. For instance, crosslinked CDHA/calcium polyacrylate biocomposites were created when TTCP and PAA interacted [534]. TTCP hydrolyzes to CDHA in aqueous solution [16], and the calcium cations that are liberated then combine with PAA to form a crosslinked network [534]. In another study, dicarboxy polyphosphazene crosslinkable with calcium cations was used as the polymeric component, and CDHA/polyphosphazene biocomposites possessing a porosity of around 65% and a compressive strength of approximately 10 MPa were created [535]. To mimic the properties of PMMA cements, PPF/β-TCP biocomposites were prepared by adding vinyl monomers to crosslink PPF. As a result, fast hardening and degradable biocomposites with low heat generation and compressive strength in the range of 1–12 MPa were prepared by varying the molecular weight of PPF, monomer, β-TCP, initiator, and NaCl as porogen [536,537]. Injectable polydimethylsiloxane/HA biocomposites [538], acrylic formulations with Sr-containing HA as filler [100], and formulations with PLGA microspheres and self-hardening CaPO₄ [539,540], along with hybrid chitosan oligosaccharide/gelatin/CaPO₄ biocomposites [541], have also been investigated.

Numerous researchers used different polymers in self-hardening CaPO₄ compositions to increase their mechanical capabilities. To stabilize the CaPO₄ paste in an aqueous solution before it sufficiently hardens and then to boost its compressive strength, for instance, gelatin was added [414,483,542]. The addition of rod-shaped fillers also improved the mechanical properties [542]. For instance, the mechanical characteristics of self-hardening TTCP + DCPD formulations were successfully enhanced by PAA and PVA, but, unfortunately, both processability and hardening time were inevitably reduced [543,544]. Similar results have been reported after the addition of sodium alginate and sodium polyacrylate [545]. Other polymers such as polyphosphazene can also be used [546,547,548]. Other examples of polymer/CaPO₄ self-hardening biocomposites can be found elsewhere [549,550]. In addition to polymers, metal wires can also be used as reinforcement to create biocomposites from the self-hardening CaPO₄ formulations [551].

Porous CaPO₄ scaffolds were created by covering PU foam with a self-hardening TTCP + DCPA formulation and calcining it at 1200 °C. Those scaffolds had interconnected macropores (~1 mm), micropores (~5 µm), and substantial porosity (~80%). To improve their mechanical properties, PLGA solution was infiltrated into the open micropores of the struts to obtain a composite construction made of biodegradable polymers and bioactive ceramics. An additional coating of 58S bioactive glass/PLGA composite was applied to the PLGA-filled struts. The resultant porous biocomposites are suitable for low-load tissue engineering scaffolds [552]. There are also more complex structures in which the PLGA macroporous phase is reinforced with a self-hardening TTCP + DCPA formulation and the entire structure is surface-coated with a CDHA layer [553]. Furthermore, a moldable CaPO₄-based self-setting composite was successfully constructed by infiltrating a wollastonite/CaPO₄ paste into a 3D plotted PLGA network. After being heated to 50 °C, the PLGA/wollastonite/CaPO₄ composite appeared to be sufficiently plastic for successful adaptation to the complex-shaped bone defects [554]. In another study, PLGA microparticles coated with α-TCP and poly(ethylene phosphate) sodium salt were mixed with castor oil and water to form oil-in-water emulsions. The resulting emulsions solidified spontaneously at room temperature in a humidified atmosphere. The PLGA particles included in the set formulation underwent hydrolysis, leading to the formation of interconnected microporosity [555].

The addition of water-soluble mannitol imparted 42–80% porosity to the self-hardening CaPO₄/chitosan biocomposites [556]. Chitosan significantly increased the mechanical strength of the overall formulations [557]. A similar approach was used by other researchers who studied the effect of adding PLGA microparticles [558,559,560,561] (which can also be filled with drugs or growth factors [562,563,564]) to self-hardening CaPO₄ formulations. Those biocomposites were implanted into cranial defects in rats and the best results were obtained with a particle content of around 30 wt% [558]. On the other hand, after 2 weeks, the biocomposites containing growth factors greatly enhanced bone contact and, after eight weeks, encouraged the creation of new bone [563]. In an in vivo rabbit femoral implant study, the self-hardening PLGA/CaPO₄ formulation showed compatibility and bioactivity as well as better osteoconductivity and degradability than the self-setting formulation consisting of CaPO₄ alone [559]. In 2018, researchers succeeded in producing bionic biocomposites with the Harvarsin microstructure from a self-setting TTCP + DCPD formulation combined with small intestinal submucosa [565]. Further details on this topic are available in the literature [566]. In addition, several types of self-hardening CaPO₄-containing biocomposites were successfully tested as inks for the 3D printing of scaffolds [486,567] (see Section 5.9 below).

Finally, the addition of CaPO₄ to self-hardening calcium sulfate (plaster of Paris) [568,569], or vice versa [570], or to calcium silicate (Portland bone cement) [571] formulations can improve their osteoconductive properties. More complex formulations, such as calcium sulfate/β-TCP/PVA/polyvinylpyrrolidone [572], are also known.

5.3. Nanosized CaPO₄-Based Formulations and Nanosized Biocomposites

Nanosized and nanophase materials are materials with particle, crystal, or grain sizes of less than 100 nm. However, before continuing, a clear distinction needs to be made between nanosized composite materials and composite materials containing nanosized particles. The former, which can be any type of composite material, is fragmented into particles with sizes smaller than 100 nm. The second consists of two or more materials, at least one of which is on the nanometer scale. Nevertheless, both possibilities combine the advantages of the two parts to offer more usefulness than just the individual materials.

When compared to large-grained materials with the same chemical composition, materials that are nanosized and nanophase have distinct mechanical and optical properties [573]. Furthermore, as compared to traditional micron-sized materials, they have different surface characteristics including more atoms, flaws, and grain boundaries on the surface, a larger surface area, and a different electrical structure. Regarding CaPO₄, the surface roughness of nanosized HA (size ~67 nm) is as high as 17 nm, compared to 10 nm for conventional submicron-sized HA (~180 nm), and the contact angle (a quantitative wetting indicator of solids by liquids) is significantly lower, at 6.1 for nanosized HA compared to conventional HA (11.51). Furthermore, the pore diameters of nanosized HA compacts are approximately 6.6 Å, which is 3–5 times smaller than the pore diameters of normal particle-sized HA compacts, which range from 19.8 to 31.0 Å [574]. Moreover, compared to microcrystalline CaPO₄, nanosized ones accelerate the production of new bone tissue by promoting osteoblast adhesion, differentiation, proliferation, osteointegration, and surface mineral deposition [575,576,577]. Finally, nanosized HA was found to induce apoptosis of leukemic P388 cells [578].

Biogenic hierarchical biocomposites based on nanosized materials make up natural bones and teeth, as nanosized plate-shaped crystals of biogenic apatite with dimensions of 50 nm × 25 nm × 1–4 nm, growing in close contact with a bioorganic fiber-rich organic matrix; thus, they generate a significant organic/inorganic interface due to the nanoscale dimensions and are arranged in a complex hierarchical structure. When calculating the interface area between the biological apatite and the organic phase based on the crystal dimensions, it becomes ~210 m²/g. This interface is believed to absorb externally applied forces to dissipate mechanical energy to thermal energy [579]. Given that the main bio-organic phase of bone is collagen, i.e., a natural polymer (Table 1), bio-composites of nanosized CaPO₄ and biodegradable polymers are clearly advantageous as bone graft materials. As mentioned before, the polymeric phase offers elasticity, whilst the CaPO₄ phase is in charge of biological activity and mechanical strength (hardness). Moreover, CaPO₄’s solubility is influenced by its carbonate content and crystal size. The more carbonate present, the more soluble the compound is. The smaller the crystals, the more soluble the compound is. To the best of the author’s knowledge, only the apatites (HA, FA, and CDHA) listed in Table 3 were accessible in nanoscale form until recently. However, in an effort to treat dental caries, nanosized MCPM [580] and DCPA [455,456,457] have recently been synthesized and used to construct biocomposites with potent ion-releasing capabilities. Probably, all known types of CaPO₄ from Table 3 can be produced in nanosized and/or nanocrystalline states, but not all of them have been prepared yet [577].

The mineralization, biocompatibility, and mechanical characteristics of biocomposites based on different (bio)polymers and nanosized CaPO₄ have all been the subject of recent studies (especially HA). Unfortunately, in most of the published papers, it remains unclear whether ‘nanosized HA’ actually represents nanosized stoichiometric HA or nanosized non-stoichiometric CDHA, and thus the two cannot be distinguished. Those studies include biocomposites of HA or CDHA with PGA [581], PLA [221,233,582,583,584,585,586,587] and its copolymers with PGA [588,589,590] and PCL [591,592], PCL [593], collagen [594,595,596,597,598,599,600,601,602], collagen + PLA [603,604,605,606,607], collagen + PVA [608], collagen + alginic acid [609,610], alginic acid [394], gelatin [611,612,613,614,615], hyaluronic acid + alkaline gelatin [616], PPF [617,618], PA [186,187,619,620,621,622,623,624], PVA [291,292,625,626,627], PVAP [295], BSA [628], poly(ethylene-acrylic) acid [629,630], chitosan [631,632,633,634,635,636,637,638] and derivatives [639], konjac glucomannan + chitosan [640,641], chitosan + gelatin [642,643,644], xanthan [645], PHEMA + PCL [646], PCL [332,371,647,648,649], PU [650,651,652], PLA + PU [653], DNA [654,655], cellulose [41,42,656,657,658,659], Ti [660,661,662], SiC [663,664], ZrO₂ [665,666,667], glass [668,669], metals and alloys [670,671], carbon quantum dots [672,673], clay [674,675], and many other biocompatible hybrid formulations [199,219,272,290,360,676,677,678,679,680,681,682,683,684,685,686,687,688,689]. A schematic drawing of DNA incorporation inside the HA structure and a TEM micrograph showing CDHA nanoparticles incorporating DNA are shown in Figure 3 [654,655].

Furthermore, any formulation can be coated with a nanosized CaPO₄ layer. In other words, nanosized HA rods + gelatin biocomposites are coated with nanosized HA [690]. Certain kinds of nanosized biocomposites seemed to be useful as carriers for the transport of drugs and growth factors [669,691,692,693], as well as potential vectors with extremely high gene loading and transfection effectiveness [694]. There is information available regarding those biocomposites’ exceptional biocompatibility [599]. The state of dispersion of nanosized particles seems to be an important parameter in regulating the mechanical characteristics of nanosized biocomposites since their high surface energy invariably causes them to agglomerate [360]. The mechanical characteristics of biocomposites comprised of PA and nano- and micron-sized HA were compared. The findings demonstrated that, as the amount of nanosized HA in the biocomposites increased, their flexural and tensile strength increased as well, but, as the amount of micron-sized HA content increased, it dropped [186]. Figure 4 displays SEM images of the mineralized collagen fibers, highlighting the intimate connection between the mineral phase and the restructured collagen fibers as well as the homogeneity of the nanosized biocomposites [695].

By means of precipitation and lyophilization, nanosized HA and porous (porosity ~85%) biocomposites consisting of collagen were created. Those biocomposites showed no pH decrease in in vitro degradation [603,604,605]. They were implanted in rabbit radii, showed high biocompatibility, and were partially resorbed after 12 weeks. From HA/chitosan nanosized rods, nanosized HA/chitosan biocomposites with enhanced mechanical stability were created [696]. Using heat-induced phase separation, nanosized HA/PLLA biocomposites with high porosity (~90%) were created [697]. Additionally, nanosized HA was utilized to create biocomposites with PAA, and the resulting nanosized crystals’ nanostructure revealed a core–shell configuration [698,699]. Other researchers discovered that the self-organization of nanodimensional HA with PAA led to the formation of liquid–crystalline nanorod hybrids that formed aligned films and showed stimuli-responsive properties [700].

Nanosized crystals of HA were found to be suitable for endosteal implantation and offered the possibility of creating advanced biocomposites for clinical applications [701]. Thus, the biocompatibility of chitosan in osteoblast cultures was significantly improved by the addition of nanosized HA [702]. Similar findings apply to nanosized HA/PA biocomposites [620]. Additional information about nanoscale biocomposites can be found in the outstanding reviews [22,703]. More specifically, there is also a review of the application of nanosized biomaterials in orthopedic surgery [704] and a review of composites and hybrid formulations containing functional CaPO₄ in nanomedicine [705]. More information on this topic can be found in other reviews [218,706,707,708,709].

To conclude, currently (May 2024), according to Scopus, there are 3110 articles with a combination of “composite”, “apatite”, and “nano” in the title, and 422 articles with “composite”, “calcium phosphate”, and “nano” in the title.

5.4. Collagen-Containing Biocomposites

A characteristic feature of bone is the nanoscale spatial orientation of bioapatite crystals and collagen macromolecules [18,19]; the crystals (about 50 nm long) are aligned parallel to collagen fibers (Figure 5a) [710] and are thought to be the source of bone’s mechanical strength. Collagen refers to a family of fibrillary proteins with a triple-helix structure of polyproline-II type. There are many types of collagens that differ in their ratios of helical to nonhelical domains, but all of them share a characteristic triple α-helix supramolecular structure that results from repeating glycine-X-Y sequences, where X and Y are typically proline and hydroxyproline, interspersed with alanine residues. Significantly, collagen type I represents 90% of the collagen present in the human body, mainly in skin, bones, tendons, and organs [711].

The mineralized fibers, composed of collagen molecules and biological apatite crystals, are approximately 300 nm long and 6 nm in diameter. The strengthening impact of nanosized bioapatite crystals in calcified tissues can be described by load transfer to inherently rigid inorganic crystals, because the collagen matrix functions as a load transfer medium, even though all of the mechanisms behind the osteogenic approach are yet unknown [14,15,19,20,21]. Furthermore, the bioapatite crystals present between the entangled fibers cross-link the fibers through mechanical interlocking or by forming bridges of calcium ions and increase the deformation resistance of the collagen fiber network [712].

Since biological apatite and collagen are the major constituents of calcified tissues of mammals, many attempts have been made to combine these two substances in biocomposites in attempts to mimic both the structure and composition of bones. The authors of a recent publication created a plot, showing that, in each passing year since 1990, the number of published works has increased exponentially (Figure 6) [713]. As seen from this figure, HA/collagen biocomposites are widely explored in the literature. Their physical forms include fibrous matrices, powders, dense materials, films, and gels. The details and references to preparation techniques of all these forms are well described in a recent review [713].

When combining CaPO₄ and collagen in the laboratory, the goal should be to create a material that closely resembles the chemical composition, microstructure, and porosity of bone, but, unfortunately, the prepared biocomposites are very different from bone (Figure 5). The introduction of an organic matrix initially, followed by the development of an inorganic reinforcement phase inside this organic matrix, is the primary characteristic of the process by which mineralized hard tissue is created in vivo [14,15,20,21]. Thus, the best technique to simulate bone is to simulate its formation. This means nucleating and growing nanosized CDHA crystals on and within collagen fibers from supersaturated solutions. Such a synthesis is called ‘biologically inspired’ [714,715,716], meaning it produces an order and environment very close to those of nature. There is a long history of using collagen and CaPO₄ (mostly apatite) in biologically inspired biocomposites for bone replacement [717,718,719,720,721,722,723,724,725,726,727,728,729], starting with the pioneering work of Banks et al. growing CDHA on reconstituted calf collagen bands in 1977 [730], followed by the first medical applications by other researchers in 1982 [731,732]. Such combinations appeared to be bioactive, osteoconductive, and osteoinductive [602,733,734,735] and, in general, artificial grafts produced from such biocomposites behaved similarly to bone and were found to be more useful in surgery than those prepared from other materials. Certainly, the data showed that CaPO₄/collagen biocomposite scaffolds were superior to artificial polymer scaffolds and CaPO₄ bioceramic scaffolds separately [736]. In addition, multi-scale simulations of apatite/collagen composites were performed [737].

CaPO₄ has been found to successfully precipitate on collagen substrates regardless of their shape or source [594,730,738,739]. Sometimes precipitation can be performed under the presence of external forces, such as an electric field [740]. However, the binding of CaPO₄ crystals to collagen depends on the degree to which collagen is denatured; the more fibrillar the collagen, the greater the binding. The first report on biocomposites produced by precipitating DCPD on collagen matrices with the help of phosphorylated amino acids was published in 1993 [717]. Moreover, CDHA crystals were shown to nucleate on the collagen fiber network when self-hardening CDHA-forming preparations (DCPD + TTCP) were combined with a collagen solution, hydrated, and hardened. This resulted in a material with worse mechanical properties than those documented for bone. Unfortunately, the prepared biocomposites did not have a bone-like nanostructure [718,741]. Using an experimental setup where Ca²⁺ and PO₄³⁻ ions diffused into collagen disks from opposite orientations, the directed development of OCP crystals on collagen was obtained [742,743]. However, those tests did not examine the mechanical properties of the biocomposites; rather, they were merely intended to mimic the in vivo apatite precipitation mechanism [744].

Traditionally, collagen/CaPO₄ biocomposites have been produced by blending or mixing CaPO₄ with collagen by means of various techniques, as well as by biomimetic precipitation [594,602,604,692,714,715,716,717,718,727,738,742,743,744,745,746,747,748,749,750,751,752,753,754]. For example, apatite/collagen bone-like biocomposites were prepared by two different methodologies: (1) dispersion of apatite in collagen aqueous suspension followed by lyophilization and (2) direct nucleation of the apatite phase on collagen fiber aggregates [716]. The biocomposites obtained by the first method resembled uncalcified natural collagen. It is evident from Figure 5b that there was no genuine contact between the apatite and collagen fibers since the crystallite size was not uniform, frequently aggregated, and dispersed randomly in the matrix. The second method allowed nanosized apatite crystals to nucleate directly on the self-assembling collagen fibers. In the latter case, CDHA and collagen showed a strong interaction (Figure 5c), which was highlighted by various analytical techniques showing that the biocomposites completely resembled the calcified natural tissue [716]. Other fabrication techniques are known as well. For example, collagen/apatite biocomposites mimicking the nanostructure of bone have been prepared using a polymer-guided liquid precursor process, in which nanosized apatite crystals are embedded into collagen fibers [750]. Much better results were obtained by other researchers, who produced multilevel hierarchical HA/collagen biocomposite scaffolds with a biomimetic bone Haversian microstructure by a combination of electrospinning, biomimetic mineralization, and rolling techniques for a bottom-up synthesis from nano-level to micro-level and then to macro-level. MicroCT scanning and scanning electron microscopy confirmed the successful preparation of the HA/collagen scaffolds composed of fully assembled microscopic fibers, while the mechanical properties of the scaffolds were double-reinforced by many newborn HA nanoparticles and chemical bonds [753]. Freeze-drying can be used as well [754]. More complex biocomposites, such as magnetite-enriched HA/collagen [755], HA/osteocalcin/collagen [487], HA/collagen/PLA [603,604,605,606,607], HA/collagen/alginic acid [609,610], HA/collagen/PVA [608,756], HA/collagen/chitosan [757], HA/collagen/acidic gelatin/basic fibroblast growth factor (b-FGF) [758], and calcium silicate/β-TCP/collagen [759], have also been developed. Recent research has shown that CaPO₄/collagen formulations have the potential to be printed by 3D printers [760].

Collagen can also be added to different self-hardening CaPO₄ formulations [718,741,761,762,763,764,765]. In order to facilitate mineral deposition, type I collagen sponges are usually soaked in a highly basic water solution containing PO₄³⁻ ions and then submerged in a solution containing Ca²⁺ ions. An alternative method involves dissolving collagen I fibers in acetic acid and mixing it with phosphoric acid. Next, neutralizing synthesis is carried out at 25 °C, with a pH of 9–10 solution, between an aqueous suspension of Ca(OH)₂ and H₃PO₄/collagen solution [714,715]. The Ca/P ratio in the reaction solution needs to be regulated in order to guarantee the quality of the finished product. Dissolving commercially available CaPO₄ in the acid is one method for doing this; another one is to add the proper ratio of Ca²⁺ and PO₄³⁻ ions to the solution and then initiate a reaction [766]. Biomimetically, CDHA crystals can be oriented and grown on dissolved collagen fibers in aqueous solution by a self-assembly mechanism [748]. Furthermore, crystallization of CDHA from aqueous solution can occur in the presence of pre-dispersed collagen [594]. More precisely, dispersing collagen in an acidic solution and then adding Ca²⁺ and PO₄³⁻ ions can cause the co-precipitation of collagen and CDHA, followed by either the pH of the solution increasing or the addition of a mixing agent. In addition, to prepare biocomposites, apatite/collagen blends were kept pressed at 40 °C and 200 MPa for several days, but biocomposites with low mechanical properties were produced [767]. Attempts have also been made to computer simulate the formation process of apatite/collagen composites [768]. It is interesting to note that apatite/collagen biocomposites appear to have some piezoelectric properties [769].

The majority of collagen/HA biocomposites are typically made by fusing HA particles, which are much smaller than microns, into the collagen matrix. This process makes it extremely challenging to create homogeneous and consistent composite grafts. Furthermore, such biocomposites have insufficient mechanical properties and, on top of that, adequate pore size cannot be achieved. In addition, unlike nanocrystalline bone apatite, microcrystalline HA may take a long time to transform into new bone tissue after implantation. Additionally, some biocomposites showed very low mechanical properties, probably due to the lack of strong interfacial bonds between the components. More to the point, in all the blend composites, the crystal size of CaPO₄ was not uniform and the crystals were frequently aggregated and dispersed randomly inside the collagen’s fibrous matrix. Therefore, only compositional similarity rather than structural similarity to genuine bone was attained. The data clearly show that bone-like chemical composition is insufficient to produce suitable grafts and that both mechanical properties and bone nanostructure need to be mimicked to function as bone at the recipient site. A grain size reduction of HA crystals to an ultrafine apatite matrix can improve its growth and osteointegration. This may improve mechanical properties, enhance biological and biochemical affinity with host bone, and facilitate osteointegration. Furthermore, it has been discovered that porosity enhances the ability of surrounding tissues to enter the pores of collagen/HA biocomposites [770,771].

Biocomposites made from bovine collagen mixed with CaPO₄ are commercially available as an alternative to bone grafts. Moreover, they might be mixed with bone marrow extracted from the iliac crest of the fracture site. Bioimplant^®, Bio-Oss Collagen^®, Boneject^®, Collagraft^®, CollapAn^®, Healos^®, Integra Mozaik^®, and LitAr^® are some examples of commercially available CaPO₄/collagen grafts that can be used in clinical practice (

Table 4. A list of known commercially produced CaPO₄-containing biocomposites and hybrid biomaterials.

Calcium Orthophosphate	Trade Name and Producer (When Available)
HA embedded or suspended in a gel	Bio-Gel HT hydroxyapatite (Bio-Rad, Hercules, CA, USA)
	Coaptite (Boston Scientific, Marlborough, MA, USA)
	Facetem (Daewoong, Seoul, Republic of Korea)
	NanoBone (Artoss, Rostock, Germany)
	Nanogel (Teknimed, L’Union, France)
	Radiesse (Merz Aesthetics, Frankfurt, Germany)
	Renú Calcium Hydroxylapatite Implant (Cytophil, East Troy, WI, USA)
HA/collagen, CDHA/collagen and/or carbonate apatite/collagen	AUGMATRIX (Wright Medical Technology, Memphis, TN, USA)
	Bioimplant (Connectbiopharm, Vienna, Austria)
	Bio-Oss Collagen (Geitslich, Wolhusen, Switzerland)
	Boneject (Koken, Tokyo, Japan)
	COL.HAP-91 (JHS Biomateriais, Sabará, Brazil)
	Collagraft (Zimmer and Collagen Corporation, Miami, FL, USA)
	CollaOss (SK Bioland, Cheonan, Korea)
	CollapAn (Intermedapatite, Moscow, Russia)
	COLLAPAT (Symatese, Vourles, France)
	DualPor collagen (OssGen, Seoul, Republic of Korea)
	G-Graft (Surgiwear, Uttar Pradesh, India)
	HAPCOL (Polystom, Moscow, Russia)
	Healos (DePuy Spine, Raynham, MA, USA)
	LitAr (LitAr, Moscow, Russia)
	nanOss Bioactive (Pioneer Surgical Technology, Marquette, MI, USA)
	Ossbone Collagen (SK Bioland, Cheonan, Korea)
	OssFill (Sewon Cellontech, Seoul, Republic of Korea)
	OssiMend (Collagen Matrix, Oakland, NJ, USA)
	OSSIX Bone (Dentsply Sirona, Charlotte, NC, USA)
	Osteomatrix (Connectbiopharm, Russia)
	OsteoTape (Impladent, Queens, NY, USA)
	ReFit (HOYA Technosurgical, Tochigi, Japan)
	RegenOss (JRI Orthopaedics, Sheffield, UK)
	RegenerOss Synthetic (Zimmer Dental, Carlsbad, CA, USA)
	SmartBone (IBI, Mezzovico, Switzerland)
	Straumann XenoFlex (Straumann, Basel, Switzerland)
	SwissBone (IBI, Mezzovico, Switzerland)
HA/sodium alginate	Bialgin (Biomed, Moscow, Russia)
rhBMP-2 + HA carrier	NOVOSIS (CGBIO, Seoul, Republic of Korea)
HA/poly-l-lactic acid	Biosteon (Biocomposites, Keele, UK)
	Fisiograft HA pasta (Ghimas, Casalecchio di Reno, Italy)
	ReOss (ReOss, Filderstadt, Germany)
	OSTEOTRANS MX (Teijin Medical Technologies, Osaka, Japan)
	SuperFIXSORB30 (Takiron, Osaka, Japan)
HA/hyaluronic acid	HArmonyCa (Allergan Aesthetics, an AbbVie Company, Dublin, UK)
HA/polyethylene	HAPEX (Gyrus, Memphis, TN, USA)
HA/glucan polysaccharide	FlexiOss (Medical Inventi, Lublin, Poland)
HA/CaSO4	BioWrist Bone Void Filler (Skeletal Kinetics, Cupertino, CA, USA)
	Bond Apatite (Augma Biomaterials, Spotswood, NJ, USA)
	Hapset (LifeCore, Chaska, MN, USA)
	PerOssal (aap Implantate, Berlin, Germany)
HA/CaSO4 powders suspended in a liquid	CERAMENT (BONESUPPORT, Lund, Sweden)
HA/glass	Bonelike (unknwn producer)
bovine bone (unsintered)	Unilab Surgibone (Unilab, Hillside, NJ, USA)
bovine bone (unsintered) + polymer	Alpha-Bio’s Graft (Alpha-Bio Tec, Petah Tikva, Israel)
	C-Graft Putty (unknwn producer)
human bone allograft	ALLOPURE (Wright Medical Technology, Memphis, TN, USA)
	AlloOss (ACE Surgical Supply, Brockton, MA, USA)
	Allosorb (Curasan, Frankfurt, Germany)
	CancellOss (Impladent, Queens, NY, USA)
	CurOss (Impladent, Queens, NY, USA)
	J Bone Block (Impladent, Queens, NY, USA)
	Maxgraft (botiss, Berlin, Germany)
	Mega-Oss (Megagen Implant, Daegu, Korea)
	NonDemin (Impladent, Queens, NY, USA)
	Osnatal (aap Implantate, Berlin, Germany)
	OsteoDemin (Impladent, Queens, NY, USA)
	Osteopure (O.S.T Laboratories, Saint-Beauzire, France)
	OsteoWrap (Curasan, Frankfurt, Germany)
	OVIS ALLO (DENTIS, Seoul, Republic of Korea)
	PentOS OI (Citagenix, Laval, QC, Canada)
	Puros (Zimmer Biomet, Stamford, CT, USA)
	RAPTOS (Citagenix, Laval, QC, Canada)
	Straumann AlloGraft (Straumann, Basel, Switzerland)
	TenFUSE (Wright Medical Technology, Memphis, TN, USA)
β-TCP/CaSO4	Fortoss vital (Biocomposites, Keele, UK)
	Genex (Biocomposites, Keele, UK)
β-TCP/poly-lactic acid	Bilok (Biocomposites, Keele, UK)
	Duosorb (SBM, Lourdes, France)
	Matryx Interference Screws (Conmed, Largo, FL, USA)
β-TCP/poly-lactic-co-glycolic acid	Evolvemer TCP30PLGA (Arctic Biomaterials, Tampere, Finland)
β-TCP/polymer	AttraX putty (NuVasive, San Diego, CA, USA)
	ExcelOs-inject (CGBIO, Seoul, Republic of Korea)
	Therigraft (Therics, Akron, OH, USA)
β-TCP/bone marrow aspirate	Induce (Skeletal Kinetics, Cupertino, CA, USA)
β-TCP/collagen	Integra Mozaik (Integra Orthobiologics, Carlsbad, CA, USA)
β-TCP/growth-factor	GEM 21S (Lynch Biologics, Franklin, TN, USA)
β-TCP/rhPDGF-BB solution	AUGMENT Bone Graft (Wright Medical Group, Memphis, TN, USA)
BCP (HA + β-TCP)/collagen	Allograft (Zimmer, Warsaw, IN, USA)
	BestAid (Maxigen Biotech, Taipei, Taiwan)
	collacone max (botiss, Berlin, Germany)
	Collagraft (Zimmer, Warsaw, IN, USA)
	Cross.Bone Matrix (Biotech Dental, Salon-de-Provence, France)
	FormaGraft (Maxigen Biotech, Taipei, Taiwan)
	Indost (Polystom, Moscow, Russia)
	MasterGraft (Medtronic Sofamor Danek, Memphis, TN, USA)
	MATRI BONE (Biom’Up, Saint-Pries, France)
	Osteon III collagen (GENOSS, Suwon, Korea)
	SynergOss (Nobil Bio Ricerche, Portacomaro, Italy)
	without trade name (MedicalGroup, Lyon, France)
BCP (HA + β-TCP)/hydrogel	4MATRIX+ (MIS Implants, Misgav, Israel)
	Eclipse (Citagenix, QC, Laval, Canada)
BCP (HA + β-TCP)/polymer	In’Oss (Biomatlante, Vigneux-de-Bretagne, France)
	Hydros (Biomatlante, Vigneux-de-Bretagne, France)
	Octane HyperPEEK (Octane, Toronto, ON, Canada)
	Osteocaf (Texas Innovative Medical Devices, Fort Worth, TX, USA)
	Osteotwin (Biomatlante, Vigneux-de-Bretagne, France)
BCP (HA + β-TCP)/chitosan	k-IBS (Bioceramed, Lisboa, Portugal)
BCP (HA + β-TCP)/fibrin	TricOS (Baxter BioScience, Maurepas, France)
BCP (HA + β-TCP)/silicon	FlexHA (Xomed, Jacksonville, FL, USA)
BCP (HA + β-TCP)/bioglass	SIGNAFUSE (Bioventus, Durham, NC, USA)
bioglass + α-TCP + β-TCP + HA + polymers	OsteoFlo NanoPutty (SurGenTec, FL, USA)
FA + BCP (HA + β-TCP)	FtAP (Polystom, Moscow, Russia)
DCPA + MgHPO4·3H2O + SiO2 + carboxymethyl cellulose	Novogro (OsteoNovus, Toledo, OH, USA)
DCPD/collagen	CopiOs Bone Void Filler (Zimmer, Warsaw, IN, USA)
DCPD + β-TCP/CaSO4	PRO-DENSE (Wright Medical Group, Memphis, TN, USA)
DCPD + β-TCP/CaSO4 + collagen	PRO-STIM (Wright Medical Group, Memphis, TN, USA)
OCP/fibrin	FibroFor (BioNova, Moscow, Russia)
OCP/collagen	Bonarc (Toyobo, Tokyo, Japan)
undisclosed CaPO4 + biologics	i-FACTOR (Cerapedics, Broomfield, CO, USA)
3D-printed Ti mesh embedded in MCPM + β-TCP + Ca pyrophosphate	OssDsign (OssDsign AB, Uppsala, Sweden)

). When treating acute fractures of long bones with defects stabilized by internal or external fixation, such materials have been compared with autografts [772,773]. The biocomposites were discovered to be osteogenic, osteoinductive, and osteoconductive; nevertheless, they are not structurally strong and necessitate the extraction of bone marrow from the patient. OCP/collagen biocomposites have undergone investigation [774] and clinical testing [775], as well.

Lyophilization was used to create open-pore (30–100 µm) collagen sponges, which were then coated with a 10 µm layer of biomimetic apatite that was precipitated from artificial bodily fluids [776]. When using fibroblast culture in vitro, researchers discovered good results. There are several possible preparation methods [777]. Injectable bone fillers have been created using collagen/HA microspheres and gel beads [778,779,780], while a bone structure was mimicked by blending carbonated apatite and collagen [781,782]. When a similar material (mineralized collagen) was implanted into rat femurs, excellent clinical results were observed after 12 weeks [783]. Collagen/HA biocomposites were prepared and mechanical performance was improved by crosslinking collagen fibers with glutaraldehyde [595,596,597]. Those biocomposites were tested in rabbits and showed good biological performance, osteoconductivity, and biodegradability. In a similar approach, HA/collagen microspheres (diameter ~5 µm) were prepared by water-in-oil emulsification and their surfaces were crosslinked with glutaraldehyde [779]. The material performed well in vitro with osteoblast culture. Porous bone graft substitutes were created by lyophilization of nanosized HA/collagen biocomposites and PLA. It was discovered that the resultant material resembled real bone at various hierarchical levels [604]. Further in vitro tests verified that osteoblasts adhered, proliferated, and migrated to this composite with acceptable quality [603]. Dopants can be added to increase biocompatibility even further. For example, Si-substituted HA/collagen biocomposites with silicon preferentially incorporated into the collagen phase have been developed to enhance bone replacement [596]. Porous (porosity ~95%, 50–100 µm interconnected pores) biocomposites of collagen (cross-linked with glutaraldehyde) and β-TCP were prepared by freeze-drying techniques followed by the sublimation of solvents. The biocomposite showed good biocompatibility when implanted in the rabbit jaw [783].

Since biocomposites of CaPO₄ with collagen were found to be able to form bioconjugates (bioconjugation is a chemical strategy to form a stable covalent link between two molecules, at least one of which is a biomolecule [784]), they are beneficial for the delivery of drugs, growth factors, and other biomolecules [610,720,765,785,786,787]. Namely, 5 weeks after implantation, HA/collagen-alginate (20 µL) supplemented with rhBMP-2 (100 µg/mL, 15 µL) demonstrated osteogenesis along the implant without the material deforming much [610]. The results of another study showed that HA/collagen biocomposites coupled with rhBMP-2 could significantly improve the formation of new bone and angiogenesis within the scaffolds as well as the proliferation and differentiation of osteoblasts [784]. Furthermore, a bilayer collagen/β-TCP implant supplemented with cartilage-inducing growth factors was developed for the repair of osteochondral defects in the trochlear groove in miniature pigs [785]. More details on this topic can be found in a recent review [786].

Although biocomposites of apatites with collagen appear to be a very hot topic of research (Figure 6 and Table 4), not many papers are devoted to biocomposites of other CaPO₄ with collagen [745,746,747,774,775,785,788,789,790,791,792]. Among them, there are ones with a very complicated composition. For example, a collagen hydrogel incorporating platelet-derived growth factor-loaded photopolymerizable zinc-based zeolitic-imidazolate-frameworks-polydopamine nanoparticles (PDGF@ZIF-8-PDA@COL hydrogel) was prepared and then perfused into PLGA-TCP scaffolds. The resulting PDGF@ZIF-8-PDA@COL/PLGA-TCP biocomposite scaffolds sustainably released PDGF and showed an excellent antibacterial activity against both the Gram-negative bacterium E. coli and the Gram-positive bacterium S. aureus. Further, the scaffolds exhibited excellent osteoinductive and osteoconductive capacity in vitro and accelerated bone formation in vivo in a rat model of a cranial defect [792].

In conclusion, currently known CaPO₄ and collagen biocomposites mimic natural bone to some extent and subsequent biological evaluations show that they do not act as permanent implants but are readily involved in bone metabolism in the form of bone remodeling [604,732]. However, the performance of such biocomposites depends on the source of the processed collagen. In an effort to enhance the mechanical characteristics of these biocomposites, numerous attempts have been made to mimic the collagen-HA interfacial behavior in actual bone using crosslinking agents such glutaraldehyde [595,597,598,738,779,783]. Unfortunately, the progress in this direction has been limited by high costs, difficulties in controlling cross-infection, poorly defined commercial sources of collagen, and a lack of suitable techniques to produce bone-like microstructures. However, since collagen is a protein that makes up the bulk of the organic bone matrix, the synthesis of CaPO₄/collagen biocomposites has been extensively studied for decades. Additional information on this hot topic can be found in the literature [602,710,713,723,793].

Finally, currently (May 2024), according to Scopus, there are 1360 articles with a combination of “collagen” and “apatite” in the title and 438 articles with “collagen” and “calcium phosphate”.

5.5. Biocomposites with Other Bio-Organic Compounds

Because the commercial sources of collagen type I are not clearly defined, the two main practical issues with this material are its expense and the challenge of adhering to a well-controlled processing method. Therefore, different substances can be used in place of collagen type I. It should be highlighted that, in addition to collagen, the bodies of humans and other mammals contain a wide variety of bio-organic substances, proteins, and biological macromolecules, all of which have substantial quantities that may be used to create biocomposites with CaPO₄ [709]. That is why this topic appears to be broad. For example, the biologically strong adhesion between teeth and the surrounding epithelial tissue (preventing bacterial invasion) is associated with the cell adhesion protein laminin [794]. To mimic this property, a biomimetic approach has been used to create laminin/apatite biocomposite layers on both titanium [795] and EVOH [796,797] surfaces. It was discovered that the more intricate laminin/DNA/apatite biocomposite layer was a successful gene delivery mechanism [798]. More details on this topic can be found in the special review [799].

Let me begin with gelatin (this is a water-soluble protein with a molecular weight from 20 to 100 kDa derived from hydrolysis of collagen type I), which is extensively used in the food industry. CaPO₄/gelatin biocomposites have been widely investigated as potential bone replacement biomaterials [196,409,410,411,412,413,414,415,432,483,484,485,611,612,613,614,615,800,801,802,803]. For example, gelatin foams have been successfully mechanically reinforced with HA and then crosslinked with carbodiimide derivatives [196]. Such foams have been shown to be good carriers for the antibiotic tetracycline [801]. In addition, electrospun gelatin methacryloyl fibrous scaffold incorporating β-TCP was developed for vital pulp therapy [804].

Since the organic–mineral interface in bones is predominantly polysaccharides [805], various biocomposites of CaPO₄ and polysaccharides, such as alginates [413,609,610,612,715,806,807,808,809], as well as chitosan [426,450,468,490,519,556,631,632,633,634,635,636,637,638,639,640,641,642,643,644,678,696,702,810,811,812,813,814,815,816,817,818,819,820,821] and chitin [416,489,822,823,824,825,826], are also very popular. For example, hydrogel-based porous HA/alginate composites have been prepared biomimetrically [715] and by lyophilization [806]. Furthermore, solution-based methods have been developed to combine chitin with HA powder homogeneously dispersed with ceramic particles [822,823]. Unfortunately, it has been difficult to achieve a homogeneous dispersion. The mechanical properties of the final biocomposites were not very good. Both tensile strength and modulus were shown to decrease with increased HA levels, likely due to poor adhesion between the filler and matrix. Microscopy showed that HA particles penetrated between the polymer chains, weakening the interaction and reducing the overall strength [822,823]. Other fabrication techniques can be found in the references, but an interesting approach is the production of HA/chitosan biocomposites from the natural CaCO₃/chitosan of crab shells by a hydrothermal process [812]. Similarly, pre-produced DCPD/chitosan biocomposites were converted into HA/chitosan biocomposites by hydrothermal procedures [813]. It was discovered that natural HA and chitosan biocomposites have outstanding osteoconductivity and good hard tissue biocompatibility, which qualifies them as framework materials for tissue engineering and artificial bone implants [811]. Research indicates that, when CaPO₄ is added to chitosan instead of chitosan alone, the result is enhanced cell adhesion, high cell proliferation, and well-diffused morphology [633,814]. More complex biocomposites, such as HA/chitosan reinforced with silk fibers [827], HA/collagen/chitosan [828], CDHA/chitosan filled with ciprofloxacin [818], chitosan/silk fibroin/nitrogen-doped carbon quantum dots/α-TCP [829], and others [830,831,832,833], have also been investigated. Besides biomedical applications, nanosized HA and chitin/chitosan biocomposites can be used for the removal of Fe(III) [834] and fluoride [835,836] ions from aqueous solutions. Further details on this topic are available in recent reviews [837,838].

Biocomposites with soluble proteins, such as CDHA/BSA, can be prepared by precipitation methods [549,839,840,841,842]. In such biocomposites, BSA does not bind strongly to solid CDHA, which is beneficial for sustained release. However, this is not the case when using the water/oil/water interfacial reaction route [190]. More recently, DNA has been incorporated into CDHA/BSA biocomposites [190,843,844,845]. In addition, unspecified CaPO₄/DNA [846,847] and biomimetic apatite/C₆₀ fullerene/Au-DNA nanosized particles [848] biocomposites have been prepared.

Additionally, CaPO₄-based biocomposites were produced by immersing hydrogels into supersaturated solutions of Ca²⁺ and PO₄³⁻ ions to precipitate CDHA (up to 70 wt% CDHA can be incorporated) [849]. Chitosan underwent a similar process, with the form of the initial chitosan hydrogel determining the 3D shape of the final biocomposites [850]. A different research team created biocomposites using self-assembled supramolecular hydrogels that underwent in situ CaPO₄ mineralization [851]. In addition, there are further experimental strategies that could be used [852,853].

There are numerous CDHA biocomposites that have been made utilizing glutamic and aspartic acids, as well as polyglutamic and polyaspartic acids [296,297,854,855,856,857]. By quantitatively incorporating these (poly)amino acids into CDHA crystals, both the coherent length of the crystal domains and the crystal dimensions were reduced. The relative amount of (poly)amino acid content in the solid phase increased with concentration in solution, reaching the maximum value of approximately 7.8 wt% for CDHA/aspartic acid and 4.3 wt% for CDHA/glutamic acid biocomposites, respectively. Small crystal sizes and large surface area due to the presence of (poly)amino acids revealed the potential of these biocomposites for hard tissue replacement [296,297,854,855,856,857]. A schematic diagram of biocomposite formation from ACP and amino acids is shown in Figure 7 [858].

Macroporous scaffolds of BCP (HA + β-TCP)/agarose with controlled complete interconnections, high porosity, fully open pores, and tailored pore size have been prepared for tissue engineering applications [859,860]. The biodegradable polymer agarose was chosen as the organic matrix as it is a biocompatible hydrogel and acts as a gelling agent, resulting in a strong gel and rapid room-temperature polymerization. A combination of cryoforming, stereolithography, and two different drying techniques produced porous scaffolds with an engineered structure. The biocompatibility of that BCP/agarose system was tested using mouse L929 fibroblasts and human SAOS-2 osteoblasts at different colony formation times [861].

Fibrin sealant is a non-cytotoxic, fully absorbable biomatrix that forms fibrin clots structured like physiological thrombi, mimicking the final stage of the natural coagulation cascade [862]. Clinical use of bone substitutes may be advanced by CaPO₄ biocomposites with fibrin sealant. The macro- and microporous structures of CaPO₄ ceramics are interwoven with a 3D network of fibrin sealant; as a result, the physical, chemical, and biological characteristics of CaPO₄ bioceramics and fibrin sealant were combined to create a biocomposite that may be used to prepare advanced bone grafts [491,863,864,865].

There are also biocomposites composed of CaPO₄ and bisphosphonates [866], silk fibroin [188,679,680,867,868,869,870,871,872], cellulose [873,874], chitosan + silk fibroin [875], chitosan + zein [876], chitosan derivatives [877], fibronectin [878], triazintrione [879], chondroitin sulfate [734,880], casein phosphopeptide [881], okra hydrocolloid [882], keratin [883], gellan gum [884] and graphene oxide/chitosan [885] hydrogels, amyloid fibrils [886], agarose [887], glycine [888], and vitamins [889]. More complicated formulations, such as sodium alginate/N,O-carboxymethyl chitosan/aldehyde hyaluronic acid/BCP, in which the amount of BCP was varied from 40% to 60% of the hydrogel mass to mimic the composition of the native bone tissues [890], are also known. Photopolymerizable formulations have also been developed [891,892]. Namely, experimental composite resins composed of 75 wt% bisphenol a-glycidyl methacrylate and 25 wt% triethylene glycol dimethacrylate were produced. Some 1 mol% trimethyl benzoyl-diphenylphosphine oxide was used as a photoinitiator and butylated hydroxytoluene was added as a polymerization inhibitor. Silica (1.5 wt%) and barium glass (65 wt%) particles were added as inorganic fillers. For remineralizing and antibacterial effects, α-TCP (10 wt%) and myristyltrimethylammonium bromide (5 wt%) were added, respectively. The authors concluded that the addition of α-TCP and myristyltrimethylammonium bromide promoted remineralizing and antibacterial properties of the resins [892].

The reader is also introduced to an intriguing method of crystallizing CDHA in poly(arylamine)/poly(styrenesulfonic acid) polymer capsules to produce empty biocomposite spheres that are micron in size. The thickness of the shell of the biocomposite spheres can range from 25 to 150 nm, contingent upon the quantity of CDHA that precipitates. Biocomposite capsules can be used as catalyst microreactors for bone mending or as medicinal agents [893].

An interesting phenomenon of fractal growth of FA/gelatin composite crystals (Figure 8) was obtained when two solutions containing Ca²⁺ and PO₄³⁻ + F⁻ ions were diffused from the opposite sides into a tube filled with gelatin gel [894,895,896,897,898]. The reasons for this phenomenon are not yet fully understood. Other types of CaPO₄-based composites based on DCPD and OCP have been similarly grown in iota carrageenan gels [899]. Nothing has been reported so far about potential biomedical applications of those unusual structural composites.

5.6. Injectable Bone Substitutes (IBSs)

IBSs are a promising option in bone regenerative medicine due to their ability to optimally fill complex bone defects in a minimally invasive manner. Thus, IBSs are considered a viable alternative to solid bone replacement materials representing a ready-to-use suspension of CaPO₄ microspheres [900,901], nanodimensional rods [902], or powders mixed with a viscous polymer solution acting as a liquid carrier phase. However, the addition of other compounds, such as calcium sulfate [903], is also thinkable. In general, IBSs appear as an opaque viscous paste with sufficient rheological properties to be injected into a bone defect using a surgical syringe with a needle. Sometimes, the IBS components may need to be mixed in an operating room and, in that case, it may be possible to produce or combine IBSs with blood, bone marrow, or platelet-rich plasma. Furthermore, IBSs can be easily produced under aseptic conditions. Their mechanical characteristics and consistent composition make them appropriate for biomedical applications [904,905]. All forms of IBSs can be separated into two primary categories: those that self-harden (Section 5.2 above) and those that do not cure. The latter are discussed here.

All types of IBSs require appropriate rheological properties to ensure in situ mineral phase bonding and cohesion combined with good cell permeability. The required viscosity level is usually created by the addition of suitable water-soluble polymers [92,194,614,615,906,907,908,909,910,911,912,913,914,915,916]. However, hydrogel-based IBSs are also known [917,918,919,920]. Some of them have the property of self-gelation [920]. Therefore, most of non-hardening CaPO₄-based IBS formulations can be considered as a viscous fluid subgroup of the CaPO₄/polymer biocomposites. Namely, IBSs containing a silanized hydroxyethylcellulose carrier and BCP (HA + β-TCP) have been described. This suspension is liquid in the pH 10–12 range, but rapidly gels at pH < 9 [907]. β-TCP can also be utilized to create injectable composites that enhance mechanical integrity [536]. Similarly, depending on the situation, polydioxanone-co-glycolide-based biocomposites reinforced with HA or β-TCP can be employed as moldable pastes or injectable viscous liquids [908]. The creation of linked pores is made possible by the emission of carbon dioxide during the post-injection crosslinking reaction. Moreover, Pickering emulsification has been used to create HA/poly(l-lactide-co-ε-caprolactone) biocomposite microparticles that can be injected without the need for molecular surfactants. Using water-dispersed HA nanosized crystals as particle emulsifiers and poly(l-lactide-co-ε-caprolactone) dichloromethane solution as the oil phase, stable injectable oil-in-water emulsions were produced [194]. CaPO₄-containing IBSs based on other (bio)organic compounds, such as gelatin [614,615], PVA [909], CMC [911], PAA [916], oligo(poly(ethylene glycol) fumarate) [910], chitosan + collagen [912], and silk fibroin + methylcellulose [913], have also been developed. Furthermore, photo-crosslinkable IBS compositions are also known [914].

Viscous IBS formulations based on BCP (60% HA + 40% β-TCP) suspended in a 2% aqueous solution of HPMC, which appeared to be fully biocompatible, resorbable, and easily adaptable to bone defects (due to initial plasticity), were developed [906,921,922,923,924,925]. It was discovered that the ideal BCP/HPMC aqueous solution ratio was 65/35 w/w. Expanding on this subject, IBSs can be loaded with cells [926,927], radiopaque components [928], microparticles [929] or functionalized with nucleic acids [930]. Self-curing formulations based on Si-HPMC hydrogels are also known [926]. A list of commercially produced CaPO₄-containing IBS formulations is given in

Table 5. A list of several commercial non-setting CaPO₄-based IBS and pastes with indication of producer, product name, composition (when available), and form [930].

Producer	Product Name	Composition	Form
Allergan Aesthetics	HArmonyCa™	HA, hyaluronic acid, 0.3% lidocaine	pre-mixed
ApaTech (Borehamwood, UK)	Actifuse™	HA, polymer and aqueous solution	pre-mixed
	Actifuse™ Shape Actifuse™ ABX	Si-substituted CaPO4 and a polymer	pre-mixed
Baxter (Deerfield, IL, USA)	TricOs Τ TricOs	BCP (60% HA, 40% β-TCP) granules and Tissucol (fibrin glue)	to be mixed
Berkeley Advanced Biomaterials (Berkeley, CA, USA)	Bi-Ostetic Putty	not disclosed	not disclosed
BioForm (San Diego, CA, USA)	Calcium hydroxylapatite implant	HA powder embedded in a mixture of glycerine, water and CMC	pre-mixed
Biomatlante (Vigneux-de-Bretagne, France)	In’Oss™	BCP granules (60% HA, 40% β-TCP; 0.08–0.2 mm) and 2% HPMC	pre-mixed
	MBCP Gel®	BCP granules (60% HA, 40% β-TCP; 0.08–0.2 mm) and 2% HPMC	pre-mixed
	Hydr’Os	BCP granules (60% HA, 40% β-TCP; micro- and nanosized particles) and saline solution	pre-mixed
Bioceramed (PT)	k-IBS®	BCP granules (75% HA, 25% β-TCP; 0.125–0.355 mm) and chitosan dissolved in PEG	pre-mixed
	n-IBS®	Aqueous suspension of HA powders	pre-mixed
Degradable solutions (Shanghai, China)	Easy graft™	β-TCP or BCP granules (0.45–l.0 mm) coated with 10 μm PLGA, N-methyl-2-pyrrolydone	to be mixed
Dentsply (USA)	Pepgen P-15® flow	HA (0.25–0.42 mm), P-15 peptide and aqueous Na hyaluronate solution	to be mixed
DePuy Spine (USA)	Healos® Fx	HA (20–30%) and collagen	to be mixed
Fluidinova (Maia, Portugal)	nanoXIM TCP	β-TCP (5 or 15%) and water	pre-mixed
	nanoXIM HA	HA (5, 15, 30 or 40%) and water	pre-mixed
Integra LifeSciences (Princeton, NJ, USA)	Mozaik Osteoconductive Scaffold	β-TCP (80%) and type 1 collagen (20%)	to be mixed
Mathys Ltd. (Bettlach, Switzerland)	Ceros® Putty/cyclOS® Putty	β-TCP granules (0.125–0.71 mm; 94%) and recombinant Na hyaluronate powder (6%)	to be mixed
Medtronic (USA)	Mastergraft®	BCP (85% HA, 15% β-TCP) and bovine collagen	to be mixed
Merz Aesthetics (GER)	RADIESSE®	HA particles suspended in a gel	pre-mixed
NuVasive (USA)	AttraX® putty	β-TCP and alkylene oxide copolymer	pre-mixed
Osartis/AAΡ (Münster, Germany)	Ostim®	Nanocrystalline HA (35%) and water (65%)	pre-mixed
Smith & Nephew (Memphis, TN, USA)	JAXTCP	β-TCP granules and an aqueous solution of 1.75% CMC and 10% glycerol	to be mixed
Stryker (Portage, MI, USA)	Calstrux™	β-TCP granules and CMC	to be mixed
Teknimed (FR)	Nanogel	HA (100–200 nm) (30%) and water (70%)	pre-mixed
Therics (USA)	Therigraft™ Putty	β-TCP granules and polymer	pre-mixed
Zimmer (USA)	Collagraft	BCP granules (65% HA, 35% β-TCP; 0.5–1.0 mm), bovine collagen and bone marrow aspirate	to be mixed

[930].

IBSs have better qualities because of their easy tissue regeneration, biocompatibility, and good rheological characteristics. The process of creating IBS biocomposites primarily enhances the system’s mechanical characteristics and gives the material resistance to fluid penetration; however, the amount of polymer that can be added to the paste limits these outcomes. Namely, depending on the kind and molecular weight of the polymer, there is a reported critical concentration (usually around 10%) above which the polymer begins to form a thick coating on the crystal clusters, preventing their bonding, causing plastic flow, and lowering the mechanical properties [549]. Moreover, a decrease in mechanical properties with increasing gel content was reported, and this was attributed to pore formation due to the dissolution of gelatin in the solution [542]. Therefore, although mechanical properties can be improved by the addition of water-soluble polymers, there still seems to be a limit to the application of CaPO₄-based IBS formulations to load-bearing zones [125]. More details on IBSs are available in the literature [567,823,904].

To conclude, currently (May 2024), according to Scopus, there are 232 articles with a combination of “injectable” and “apatite” in the title and 370 articles with “injectable” and “calcium phosphate”.

5.7. Biocomposites with Glasses, Inorganic Compounds, Metals, and Carbon

To overcome the problem of low mechanical properties of CaPO₄ bioceramics on the one side and induce a positive tissue response of various inorganic materials, glasses, and metals on the other side, suitable biocomposites have also been developed. Such biocomposites can be prepared by common ceramic processing techniques, such as post-mixing heat treatment [931,932,933], powder slurry coating [934] and metal-sol mixing [935]. For example, HA has been combined with Bioglass^® (Novabone Products, Alachua, FL, USA) [936,937] and other glasses [938,939] to form glass–ceramic biocomposites. The same is valid for TCP [940]. Regarding shape, the reinforcing materials for CaPO₄ include particles [941,942], platelets [943,944], whiskers [580,664,945,946,947], and fibers [948,949,950,951,952], regarding their sizes, they can be of nano- (Section 5.3 above), micro-, and macro-dimensions, and certainly they can be of diverse chemical composition. Namely, zirconia and PSZ [931,932,933,934,945,953,954,955,956,957,958,959,960,961], magnesia [962], alumina [941,944,963,964,965,966], alumina + magnesia [967], titania [968,969,970,971], alumina + titania [972], iron oxides [973,974,975,976] (Figure 9), other oxides and mixtures thereof [947,977,978,979,980,981,982], silica and/or glass [983,984,985,986,987,988,989,990,991,992,993,994], titania + bioactive glass [995], wollastonite [996,997,998], mullite [999,1000,1001,1002], natural aluminosilicates [1003,1004], nitrates [1005], various metals and alloys [185,203,204,205,551,670,671,950,1006,1007,1008,1009,1010,1011,1012,1013,1014,1015,1016,1017,1018,1019,1020,1021,1022,1023,1024,1025,1026,1027,1028,1029,1030,1031,1032,1033,1034,1035,1036,1037,1038,1039,1040,1041], calcium sulfate [1042,1043,1044,1045,1046,1047], calcium carbonate [1048,1049], silicon carbide [663,664,946], barium titanate [1050,1051,1052], zeolites [1053], boron nitride [1054,1055], zirconium nitride [1056], carbon [1057,1058], strontium hexaferrite [1059], and some other materials [1060,1061,1062] have been added to CaPO₄ to create biocomposites and improve reliability. Among them, Fe₃O₄/HA formulations have both photocatalytic [973,974] and magnetic properties [399,975,976]. Other magnetic additives to CaPO₄-containing formulations comprise ferrites CoFe₂O₄, NiFe₂O₄, CuFe₂O₄, oxides CoO, NiO, and Gd₂O₃, as well as metallic Nd, Gd, and Sm [1063]. The magnetic properties of CaPO₄-containing composites and hybrid formulations are widely used for anti-cancer treatments [1064,1065]. Some formulations can possess important electrical properties [1066]. More complex biocomposites, such as HA/Al₂O₃/carbon nanotubes [1067], PMMA/HA/ZnFe₂O₄/ZnO [530], cellulose acetate/HA/bioglass/ZrO₂ [1068], and anorthite/β-TCP/CaMgP₂O₇ [1069] formulations, have also been developed. The ability of biocomposites to form a stable interface with bone is lower than that of CaPO₄ bioceramics alone because many inorganic materials are either not bioresorbable, not bioinert, or significantly less bioactive than CaPO₄. Typically, a significant amount of reinforcing phase is required to achieve the desired properties. These formulations are sometimes called bioinert/bioactive composites because they contain bioinert components [985]. The ideal reinforcement is one that provides mechanical integrity at low loads without reducing the bioactivity of CaPO₄. CaPO₄/zirconia biocomposites (over 420 publications according to Scopus) seem to be the most popular research topic. The main drawback of PSZ-reinforced HA is the degradation of zirconia in wet environments [945,954,955,957].

There are various varieties of CaPO₄/glass biocomposites. The first, also called bioactive glass–ceramics, represent parent glasses composed of ~34% by weight of 50–100 nm sized oxy-FAP (Ca₁₀(PO₄)₆(O,F)₂) and β-wollastonite (CaO·SiO₂) crystals in a heat-treated MgO-CaO-SiO₂ glassy matrix, resulting in dense and homogeneous biocomposites [995,996,997,998]. In addition, there are A-W glass–ceramics that appear to be aggregates of small apatite particles effectively reinforced with wollastonite. A-W glass–ceramics have the highest flexural strength, fracture toughness, and Young’s modulus among bioactive glasses and glass–ceramics, making them ideal for applications involving significant compressive loads, such as iliac crest replacement and artificial vertebrae. A-W glass–ceramics combines high bioactivity with favorable mechanical properties [1070]. β-TCP/wollastonite [1071,1072,1073], β-TCP/wollastonite/diopside [1074], and OCP/silica [1075] biocomposites are also known. More complex formulations have also been developed. For example, A-W/HA/silk fibroin biocomposites are designed to match the mechanical strength of human cortical bone, provide good biological activity, and can be used for many orthopedic applications [1076]. Other examples include wollastonite-reinforced HA/Ca polycarboxylate [1077], glass-reinforced HA/polyacrylate [1078], and photocurable 3D-printed mesoporous bioactive glass/TCP [1079] biocomposites, as well as calcium phosphate silicate/wollastonite formulations containing collagen [1080] and gelatin [1081].

CaPO₄/glass biocomposites can be prepared by simple sintering of suitable CaPO₄/glass powder mixtures [1082,1083,1084,1085,1086,1087,1088]. If sintering is performed below 1000 °C, CaPO₄ does not react with bioactive glass [1083], or this reaction is limited [1084]. In a different method, CaPO₄ bioceramics were treated with tiny amounts of bioactive glass to enhance densification and/or mechanical characteristics [23]. Furthermore, biocomposites can be sintered from HA and silica powders [985]. On the one hand, the synergistic combination of two bioactive phases can modulate the dissolution of the final system and enhance its biological response, while, on the other hand, the glass acts as a sintering aid aimed at densifying the composite and increasing its mechanical strength [1085]. Under both in vitro and in vivo settings, bioactive glass–ceramics often retain more strength than HA bioceramics for an extended period of time [987]. Additionally, the creation of biocomposites by the addition of bioactive glasses to CaPO₄ has positive effects on cell adhesion, viability, and proliferation if compared to pure CaPO₄ bioceramics [1086]. More details on this topic can be found in the following reviews [1085,1086].

Regarding biocomposites of CaPO₄ with carbon, there are two major directions: bone char [1089,1090] and those with nanodimensional carbon [1091,1092,1093,1094,1095,1096,1097,1098,1099,1100,1101,1102,1103,1104,1105,1106,1107,1108,1109,1110,1111,1112]. Bone char is produced by high-temperature pyrolysis of bones (by-products of the agrifood industry) in an inert atmosphere, which converts the bioorganic compounds (mainly, collagen) into an inorganic graphitic carbon. Differently from standard biochar of plant origin, bone char also contains CaPO₄ (commonly HA). The combination of CaPO₄ and graphitic carbon makes bone char a unique material, with different possible uses; however, up until now, it has been without biomedical applications [1089,1090].

It is very difficult to prepare homogeneous mixtures of CaPO₄ and carbon nanotubes because their shapes are very different: “one can imagine something similar to achieving a homogeneous mixture of peas and spaghetti” ([1091], p. 7). Similar shape incompatibility problems are valid for CaPO₄/graphene mixtures. Nevertheless, different strategies can be used to prepare such biocomposites [1092]. For example, apatite can be chemically synthesized using carbon nanotubes with carboxyl functional groups as a matrix [179,180,181,182,1093,1094,1095,1096,1097]. The physicochemical properties of those biocomposites have shown that CDHA nucleation is initiated via carboxyl groups [179]. Hot pressing [1098], compression [1099], plasma spraying [1100], laser surface alloying [1101,1102,1103], spark plasma sintering [1104], shear mixing [1105], sol-gel [1106], and deposition [1107] techniques can also be applied. It is recommended that carbon nanotube/CaPO₄ biocomposites be sintered in a deoxidizing environment because carbon oxidizes at high temperatures [1108]. Biocomposites of CaPO₄ with carbon nanotube oxides [1109] and/or carbon nanotubes with oxygen-containing functional groups [1110] are also known. Regarding the mechanical properties, nanoindentation tests on carbon nanotube-reinforced HA composite coatings were performed to get the numerical values for hardness and Young’s modulus [1111]. The results revealed that the higher the nanotube loading, the better the mechanical properties. That is, at 20 wt% loading, hardness and Young’s modulus increased by 43% and 21%, respectively, compared to single-phase HA coating. Scratch tests showed that alloyed HA biocomposite coatings exhibited improved wear resistance and lower coefficients of friction when the number of carbon nanotubes in the precursor material powder was increased [1102]. Additionally, examinations of the biocomposite coatings’ elastic modulus and hardness revealed that the quantity of carbon nanotubes also had an impact on the mechanical properties [1101]. Compression tests on bulk HA/carbon nanotube biocomposites were conducted by another research group, which reported an improvement in strength compared to single-phase HA [1093]. The maximum compressive strength of any material, however, was only 102 MPa, which is considerably less than the usual values for high-density HA but comparable to cortical bone [1091]. Hybrid biocomposites with more intricate formulations, including poly-l-lysine/HA/carbon nanotube [1111] and multi-walled carbon nanotube/HA/carbon fiber/PEEK [1112], have also been created. Carbon nanotubes, on the other hand, are an extremely stable substance that cannot be biodegraded or reabsorbed. As a result, the nanotubes may enter the human body through the biocomposite matrix during in vivo bioresorption and result in unknown health issues. This issue needs to be resolved.

To conclude the carbon issue, carbon fibers of very small sizes [1113,1114,1115], as well as diamonds [1116,1117], graphene and graphene oxide [171,177,178,197,689,885,1118,1119], fullerenes [848] and their derivatives [498,1120], carbon quantum dots [672,673], and other allotropes of the nanosized carbons were used to reinforce the CaPO₄ bioceramics [1092]. More complicated multicomponent formulations, such as HA nanobelts/carbon nanotubes constructed into carbon fiber/DCPA/epoxy biocomposites, have been developed as well [1121]. Additional details on CaPO₄/carbon biocomposites and hybrid formulations are available in the excellent recent review [1092].

Additionally, many CaPO₄ biocomposites containing metals and alloys have been created [183,203,204,205,551,670,671,725,1006,1007,1008,1009,1010,1011,1012,1013,1014,1015,1016,1017,1018,1019,1020,1021,1022,1023,1024,1025,1026,1027,1028,1029,1030,1031,1032,1033,1034,1035,1036,1037,1038,1039,1040,1041]. Numerous fabrication techniques have been used to produce them. For example, HA-based biocomposites reinforced with 20% volume Ti particles were produced by hot pressing [1008]. A combination of cold pressing followed by sintering has also been used [1033]. Furthermore, CaPO₄ and metal biocomposites can be prepared by a powder metallurgical process [1009,1011,1021,1022], combinations of direct ink writing with a liquid pressure infiltration [203], mechanical compaction and sintering [671,1017,1019], and upward friction stir processing [1032], as well as by additive manufacturing [204,205]. Silver ions integrated into HA particles can be reduced by γ-irradiation to create Ag nanoparticle/HA biocomposites [670]. More complicated formulations, such as Ag/Ag₃PO₄/BCP (HA + β-TCP) [1035], Ti6Al4V/HA/TiB₂ [1030], HA/i-Al₆₄Cu₂₃Fe₁₃ quasicrystals [1029], and CaSiO₃/HA reinforced with a Ti6Al4V alloy [1039], were prepared as well.

In the case of metals and alloys as the matrixes, Mg–3Zn–Ca/1% β-TCP biocomposites were produced by a melt shearing technology combined with a high-pressure die-casting process. The results showed that mechanical properties were not satisfactory, probably due to the poor interfacial bonding between the phases. The authors suggested that hot extrusion and heat treatment might improve the mechanical properties [1016]. More to the point, metal matrix composite with pure Zn as a matrix and HA powder as reinforcement were prepared by a spark plasma sintering [1028], while those with pure Mg as a matrix and HA-Al₂O₃ HA-TiO₂@Al₂O₃ as reinforcements were prepared by a stir casting [1031] techniques. Additive manufacturing based on 3D printing with different content levels of powder and polymer parts, followed by washing in acetone and further sintering, was used to produce Fe-Cu-HA biocomposites [1037]. A friction stir processing technique was used as well [1020,1024]. Regarding metals and alloys with relatively low melting points, CaPO₄-containing composites and hybrid formulations with them can be prepared by infiltration of those molten metals through CaPO₄ scaffolds [203,1015,1018,1023]. A melt extrusion approach was also tested [1014].

It has been observed that, at high temperatures, the presence of Ti metallic phase accelerates both the dehydration and breakdown of HA to mixes of β-TCP and TTCP [1008,1009,1019,1122] or β-TCP and calcium titanates [661,1011,1022,1122]. Compared to pure HA bioceramics produced under the same conditions, HA/Ti biocomposites have better fracture toughness, flexural strength, and fracture work, making them suitable for biomedical applications [1017]. However, their mechanical properties did not seem sufficient for the use of HA/Ti biocomposites in load-bearing applications. Fortunately, histological evaluation revealed that HA/Ti biocomposites were partially integrated with new bone tissue after 3 weeks, and fully osteointegrated in vivo after 12 weeks [1008]. Similar findings were previously obtained for HA bioceramics reinforced by the addition of Ag particles (5–30% volume) followed by sintering of HA/Ag powder compacts [1006,1007]. Furthermore, the addition of Ag imparts antimicrobial activity [670,1013,1035]. Other studies on CaPO₄/Ti biocomposites are available elsewhere [1010,1012,1022]. In all aforementioned studies, CaPO₄ was used as the matrix, but the reverse is also possible: to investigate the effect of 100 nm sized β-TCP spherical particles on the microstructure of the alloy, biocomposites were prepared using an Mg3Zn0.8Zr alloy as a matrix and β-TCP particles as reinforcement [1123].

To conclude, CaPO₄/metal biocomposites are commercially produced, for example, OssDsign (Table 4).

5.8. Biocomposites from CaPO₄ Only

Now, biocomposites composed only of CaPO₄ should be discussed. The earliest paper on this sub-topic was published in 1976 [70]. First, all types of biphasic (BCP), triphasic, and polyphasic CaPO₄ formulations should be mentioned [436]. For example, in the 1980s, BCPs were called “TCP ceramics complexed with HA” [1124]. Even today, such multiphasic formulations are sometimes referred to as nanocomposites [1125,1126] or simply composites [1127]. Furthermore, fluorinated HA (chemical formula Ca₁₀(PO₄)₆(OH)_2−xF_x, where 0 < x < 2) can also be listed as a composite material [1128], although it is debatable whether the term ‘composite’ can be applied to such a system. Biocomposites of 70% HA-powder + 30% HA-whiskers should be mentioned and discussed instead. They were made using hot pressing, hot isostatic pressing, and pressureless sintering. Without sacrificing bioactivity or biocompatibility, those biocomposites were shown to have increased toughness, reaching the lower bound of bone fracture toughness [1129,1130]. Another illustration is a dual HA biocomposite, which combines two different porosity HA materials: one with 0% porosity for load carrying and the other with 75% porosity for bone ingrowth. The dual HA biocomposite was discovered to be a viable substitute for iliac bone grafting when used as an implant material for interspinous fusion, eliminating the disadvantages associated with autograft harvesting [1131]. In addition, two-component suspensions of HA nanoparticles with spherical (S-HA, as a matrix) and fiber (F-HA, as a reinforcement) morphologies were prepared and deposited. The results revealed that long crystalline F-HA particles efficiently reinforced the microstructure of the composite coatings; however, coarse pores were formed by the stacking of F-HA particles. Nevertheless, finer particles of poorly crystalline S-HA efficiently filled these coarse pores, providing the desired corrosion resistance [1132]. Additionally, HA-whisker-strengthened HA scaffolds have been prepared by 3D printing, followed by BCP (HA + β-TCP) coating [1133]. Furthermore, self-setting DCPD forming formulations were reinforced by the addition of CDHA/DCPA fibers [495].

A biodegradable biocomposite porous scaffold composed of a β-TCP matrix and nanosized HA fibers was developed and investigated for load-bearing bone tissue engineering; nanosized HA fibers were prepared by biomimetic deposition method and their incorporation significantly improved the mechanical properties of the scaffold, achieving a compressive strength of 9.87 MPa, comparable to the high-end value (2–10 MPa) of cancellous bone [1134]. Furthermore, HA and β-TCP powders were wet mixed, the powder combination was compacted, and the mixture was calcined to create HA/β-TCP biocomposites with varying β-TCP contents (10, 20 and 30 wt%) [1135]. Additionally, a CDHA/α-TCP biocomposite was prepared by α-TCP addition into a self-setting apatite-forming cement [1136]. Similarly, DCPA/ACP biocomposites were prepared by mixing DCPA and ACP powders followed by setting [1137]; however, it was not the case when carbonate apatite was added into a self-setting α-TCP [516]. Finally, it is interesting to mention the successful reinforcement of carbonate apatite porous blocks with newly prepared carbonate apatite crystals (i.e., the same compound) [1138]. First, calcium salts were added to the micropores of the carbonate apatite blocks. Then, in the second step, the calcium salts were carbonized by exposure to carbon dioxide gas to form calcite in the micropores of the carbonate apatite blocks. In the third step, the blocks were immersed in a Na₂HPO₄ solution. This converted calcite in the micropores of the carbonate apatite block into carbonate apatite and the newly formed carbonate apatite crystals became entangled with the crystals of the existing carbonate apatite block. The bonding of the newly formed carbonate apatite crystals with the crystals of the existing carbonate apatite block increased the mechanical strength of the block by 1.5 times compared to pre-treatment [1138].

5.9. Inks for 3D Printing

According to Wikipedia, “3D printing or additive manufacturing is the construction of a three-dimensional object from a CAD model or a digital 3D model. It can be done in a variety of processes in which material is deposited, joined or solidified under computer control, with the material being added together (such as plastics, liquids or powder grains being fused), typically layer by layer” [1139]. For the bone grafting cases, 3D printing is used to produce bioresorbable porous grafts of the patient-specific shape and dimensions according to the bones being replaced wholly or partially. This resulted in the development of CaPO₄-containing composites and hybrid formulations, which could be used as inks for 3D printing. In most cases, such formulations represent viscous liquids or pastes, which can be both self-hardened and produce solid constructs by either drying or after treatment. The latter procedure can be performed by using both physical treatments (such as light for light-crosslinking formulations), temperature treatments (freezing or sintering), and additional chemical reagents. Therefore, CaPO₄-containing inks for 3D printing could also be considered as formulations belonging to either self-hardening biocomposites (Section 5.2) or injectable bone substitutes (Section 5.6); however, since the viscous inks are not designed for in vivo injections, biologically incompatible compounds (such as organic solvents) might be used in their composition because they will be eliminated afterward during either drying or later treatments. Several examples of typical CaPO₄-containing formulations are described below [1133,1140,1141,1142,1143,1144,1145,1146].

For example, HA/PCL inks were prepared by PCL dissolving in a ternary solvent prepared by mixing dichloromethane with the surfactant 2-butoxyethanol and plasticizer dibutyl phthalate in a volumetric ratio of 5:3:1, followed by addition of a HA powder through a 30-micron mesh sieve [1140]. The influence of important variables, such as the polymer type, concentration, solvent, additives, and ceramic particle characteristics, on the printability, shape fidelity, and mechanical properties of the 3D printed scaffolds was investigated. By systematically varying the composition of HA:PCL inks, the authors demonstrated their non-Newtonian flow behavior and identified printable ink formulations based on their rheological properties. The weight ratio of HA in the PCL polymer was found to impact the compressive moduli and toughness of the prepared scaffolds [1140]. Similar results were obtained in another study, in which HA/PCL/polyethylene oxide inks both with and without adding vancomycin were prepared by a similar technique [1141]. To prepare inks, aqueous PVA solutions might be used instead of PCL solutions in organic solvents [1133].

In another study, light-curable gelatin methacrylate/HA inks were prepared by mixing 1% w/v spray-dried fine HA powders with 10% w/v gelatin methacrylate. To do this, photo-crosslinking photoinitiator Irgacure 2959 powder was dissolved in phosphate buffer saline at 70 °C for 60 min (the final concentration was 0.3% (w/v)). Then, the desired amount of gelatin methacrylate was dissolved in the Irgacure 2959 solution at 37 °C. Finally, to produce inks for 3D printing, a suspension of HA in phosphate buffer saline was prepared and applied to the gelatin methacrylate and Irgacure 2959 solution [1142].

In one more study, silk fibroin (0%, 6%, 8%, 10%, 12%), gelatin (10%), and nanodimensional HA powders (3%) were weighed according to different mass fractions and dissolved in deionized water to prepare five formulations of composite hydrogel precursors, which were used as 3D printing inks and stored in a refrigerator at 4 °C. The printing was performed at 5 °C. Subsequently, the samples were frozen for 12 h in a cryogenic storage tank at −20 °C, then taken out and thawed at room temperature for 2 h, repeating the above process 3–4 times. Finally, the samples were soaked in absolute ethanol for storage [1143].

Additionally, several formulations of HA-reinforced alginate–chitosan-based printable hydrogel inks were developed in one more study [1144]. To do this, the authors prepared an aqueous suspension of fine HA powders of different concentrations (0.1, 0.2, and 0.4 w/v %). Then, a fixed concentration of 5 w/v % alginate was dissolved in the HA suspensions through magnetic stirring. Afterward, chitosan (1 and 2 w/v %) was mixed and stirred for a few hours to swell into the alginate solution. Acetic acid was added dropwise (12–14 drops per 10 mL) in the solution with constant stirring. After printing the hydrogel inks, a 10% aqueous solution of CaCl₂ was used to cross-link the hydrogel. Afterward, the printed constructs were washed thoroughly with deionized water, followed by drying at 40 °C and storage in a vacuum desiccator [1144].

Finally, HA/PEG diacrylate hydrogel inks should be mentioned [1145]. These inks were prepared by dispersing three different concentrations of HA (1, 2, and 5 wt%) and 20 w% PEG diacrylate hydrogel (700 kDa) in distilled water. The maximum concentration of HA was set at 5 wt% because it did not block the dispensing nozzle. Then, 0.1 wt% of a radical photoinitiator for the UV curing Irgacure 2959 was added and stirred to make homogenous solutions. Afterward, 25 wt% of a non-ionic copolymer surfactant Pluronic F127 was added, and the hydrogel inks were stored in the fridge for two days for the complete dissolution of Pluronic F127. Finally, the inks were loaded into 10 mL cartridges and allowed to equilibrate to room temperature for one hour before printing to initiate the physical gelation of Pluronic F127 [1145].

It is also possible to produce porous bioceramic scaffolds from the CaPO₄-containing biocomposite inks. For example, different ink formulations consisting of Mg- and Na-doped carbonated CDHA, β-TCP, and a non-ionic copolymer surfactant Pluronic F127 were prepared and used to print porous scaffolds. Afterward, the scaffolds were first placed at room temperature for 24 h to dry, followed by sintering at 1100 °C for 2 h and cooling [1146].

To finalize this section, one should note that due to the rapid development of additive manufacturing techniques, presumably, by means of the subsequent modifications, any type of CaPO₄-containing biocomposite and hybrid formulation might be 3D printed [1147].

5.10. Functionally Graded Biocomposites

In numerous instances, a homogeneous filler distribution within the matrix is necessary [365], but there are certain formulations where this criterion is not applicable. For instance, functionally graded materials (FGMs) denote materials where the structure and/or composition gradually change in one or more directions, leading to correlated changes in the material properties. The primary characteristic of compositional FGMs is an almost continuously graded composition, which yields two distinct properties at either end of the structure. These composite materials can be tailored for particular functions and applications. Various methods, such as bulk (particle processing), pre-mold, additive, and melt processing, are employed for the production of FGMs.

Bone is a biologically created composite that exhibits varying degrees of density: from very dense and hard cortical bone to porous and foamy trabecular bone structures. Typically, cortical bone forms the outer part of long bones, and its density decreases towards the center where the trabecular bone is situated. Trabecular bone is porous, and its spaces are filled with marrow [14,15]. Therefore, bone is a naturally occurring functionally graded composite.

The concept of FGM is becoming increasingly important in the field of biomaterial design and research. Numerous studies have been conducted, for instance, on the synthesis of porous CaPO₄ bioceramics to imitate the natural porosity of bones [1148,1149,1150]. This serves as a structural approach to the development of FGMs, in addition to compositional approaches. For instance, functionally graded composite cranial implants have been produced that are made of polylactic acid, carbonate apatite, and CaCO₃ [326,327]. Additionally, HA/Ti biocomposite implants, which are functionally graded, have been prepared utilizing powder metallurgy. These implants provide biocompatible HA on the tissue side and mechanically robust Ti on the outer side [201,202,1151,1152]. The longitudinally-graded structure includes more Ti in the upper part and more HA in the lower part. In the Section 5.9, Ti bears direct occlusal forces and provides necessary mechanical performance, whereas, in the Section 5.11 embedded in bone, HA imparts bioactivity and osteoconductivity to the material [1151]. The optimal sintering conditions for Ti and HA differ significantly, creating complications in producing HA/Ti functionally graded biocomposites. Therefore, sintering conditions for the HA/Ti mixtures must be balanced. The anticipated characteristics of the implant are illustrated in Figure 10 [1152]. These biocomposites can either be symmetric [1153] or asymmetric [1154].

Functionally graded CaPO₄ coatings with different concentrations of silver were applied on Ti substrates using ion beam-assisted deposition. Cross-sectional analysis revealed a reduction in crystallinity and the dispersion of nanosized silver particles (10–50 nm) from the coating/substrate interface toward the top surface [1155]. Plasma spraying was employed to prepare compositionally graded HA/Ti biocomposite coatings [1156]. Furthermore, laser cladding [1157] and combinatorial matrix-assisted pulsed laser evaporation techniques [1158] have been utilized to deposit graded carbon nanotubes/HA, Sr-substituted HA, and zoledronate-modified HA deposits, respectively. Additionally, functionally graded HA/PMMA biocomposites were developed by dispersing precipitable HA in PMMA viscous liquids using centrifugation to prevent stress convergence at the interface. Technical term abbreviations were explained upon first use. The stress–strain curve of this biocomposite indicates sufficient strength for utilization in biomedical applications, alongside decreased brittleness and fragility [528]. Compositionally graded biocomposite scaffolds made of collagen and nanoscale hydroxyapatite can be produced through in situ diffusion techniques. Chemical and microstructural analyses showed that there is a Ca/P ratio gradient throughout the scaffold template’s width, leading to the development of Ca-rich and Ca-deficient sides of the scaffold. The Ca-rich side has low porosity and nanosized HA crystal aggregates, while the Ca-deficient side has high porosity and nanosized HA crystals integrated with collagen fibers to create a porous network structure [1159]. A biocomposite membrane with a gradient structure was created using a layer-by-layer casting method [606]. The three-layer structure consists of a porous membrane containing 8% nanosized carbonate apatite/collagen/PLGA on one side, a nonporous membrane made of pure PLGA on the opposite side, and a middle layer with a transition zone containing 4% nanosized carbonate apatite/collagen/PLGA. A similarly graded biocomposite made of PLGA/nano-HA/lauric acid was also manufactured [1160]. Functionally graded nonwoven fabrics made of PCL and containing nanosized particles of β-TCP were produced using a hybrid twin-screw extrusion/electrospinning method [1161]. Biocomposites containing functionally graded HA/silk fibroin were developed through pulsed current sintering [1162]. Additionally, HA/glass FGM layers were applied onto a Ti6Al4V substrate with the aim of creating a strong bond between the FGM laminate and the substrate [1163,1164].

Functionally graded β-TCP/FA biocomposites combine the biostability of FA with the bioresorbable properties of β-TCP. A multilayered structure, each layer being 1 mm thick, of β-TCP/FA biocomposites with varying molar ratios was produced resulting in the formation of an FGM (Figure 11). After implantation, functionally gradient porosity allowing for bone ingrowth would result from preferential dissolution of the β-TCP phase [1165]. Additionally, the same multilayered technique was used to produce HA/alumina/zirconia biocomposites with a composition gradient [1166] and 316L stainless steel/β-TCP [1167]. Two additional materials were synthesized—fluoridated HA with a fluoride gradient [1168] and carbonated HA with a carbonate gradient [1169]. Furthermore, HA/zirconia-graded biocomposites were developed to enhance HA mechanical characteristics while preserving its ability to attach to bone [958]. Titania and HA were identified as a suitable combination for FGM due to their ability to provide both a gradient of bioactivity and robust mechanical strength [1170]. Additionally, researchers have developed graded biocomposite structures of HA/CaCO₃ for improved bone ingrowth [1171]. Furthermore, functionally graded biocomposite skull implants made of polylactides, carbonateapatite, and CaCO₃ have been developed [326,327]. The research in this field shows promise, but the mechanical properties of functionally graded biocomposites do not yet match those of bones [1172].

5.11. Biosensors

Biosensors detect analytes by combining biological and physicochemical detection elements. They are made up of three parts: a biological element that is sensitive, a transducer or detector element that transforms the signal from the interaction of the analyte and the biological element into an electrical signal (current or voltage), and related electronics that present the results in an easy-to-read manner [1173].

CaPO₄ surfaces have an excellent adsorption capacity for functional biomolecules such as proteins, albumins, DNAs, and other chemical species. Therefore, CaPO₄-based composites and hybrid formulations can be applied to biosensor fabrication [1111,1174,1175,1176,1177,1178,1179,1180,1181,1182,1183,1184]. For instance, the synthesis of hybrid nanoscale particles including poly-l-lysine, HA, and carbon nanotubes was explained, and a generic immunosensing platform design method was suggested, predicated on the adsorption of antibodies onto those biocomposites [1111]. In another paper, a hybrid material consisting of gold nanosized particles mounted on nanosized HA was used to design an interface for antibody-bound piezoelectric immunosensors. The developed sensing interface appears to have advantages over the use of nanosized HA or gold alone, such as immobilization without activation and higher antigen-binding activity of the antibody [1175]. For the detection of phenolic compounds, a novel tyrosinase biosensor based on nanosized HA/chitosan composites has been developed [1178]. A Bi₂Zr₂O₇/HA composite sensor showed outstanding selectivity and performance toward sensing lead nitrate and dextrose [1183]. Due to the recent coronavirus disease-19 (COVID-19) pandemic caused by the SARS-CoV-2 virus, it is important to stress that a HA–lanthanum strontium cobalt ferrite (HA–LSCF) composite showed a good response on a screen-printed carbon electrode electrochemical aptasensor to detect the SARS-CoV-2 spike RBD protein. The authors concluded that the developed HA–LSCF-based biosensor could be used as an alternative method for detecting the COVID-19 pandemic [1184]. Finally, it may also become possible to harness the power generated by DCPD/polymer composite-based battery devices [460,461]. More details on CaPO₄-based composites and hybrid formulations as components of various biosensors can be found elsewhere [1111,1174,1175,1176,1177,1178,1179,1180,1181,1182,1183,1184].

To conclude, currently (May 2024), according to Scopus, there are 85 articles with a combination of “sensor” and “apatite” in the title and 46 articles with “sensor” and “calcium phosphate”.

6. Interactions among the Phases in CaPO₄-Based Biocomposites

One important factor that needs to be carefully examined is whether CaPO₄ interacts with other phases in composites and hybrid formulations. This interaction arises exclusively at the interface, a small region where significant alterations in the chemical composition of the constituent materials culminate in their bonding and potential load transfer. In general, interactions between phases in composites can occur through mechanical means, such as the radially compressive forces that the matrix exerts on the filler particles (for example, during cooling due to thermal shrinkage), or chemical means, where the reactivity of the filler with the matrix is of vital importance. In the latter case, it is crucial to differentiate between physical interactions and chemical bonds [268]. According to Wypych, physical interactions are typically temporary and involve hydrogen bonds and van der Waals forces. In contrast, chemical bonds are stronger and more durable as they involve the formation of covalent bonds [1185]. Consequently, composite materials’ strength is increased by preferred chemical interfacial bonding between phases. The interfacial bond’s size between the phases determines the degree of stress transferred from the weaker matrix to the stronger fibers. However, although a bond between the matrix and reinforcement is necessary for stress transfer, it is insufficient to prevent hardening mechanisms, such as phase segregation and fiber withdrawal [1091].

The precise modes of binding between the bone mineral (bioapatite) and bioorganic material (collagen)—which have a significant influence on the mechanical properties of bone—are still a contentious topic. Bone minerals do not bind directly to collagen; instead, they bind through non-collagen proteins, constituting approximately 3% of bone (Table 1), which offer active sites for biomineralization and cell adhesion [22]. The interfacial binding forces in bones are primarily ionic and hydrogen bonds, along with hydrophobic interactions which are responsible for its complex behavior [337]. In contrast, some argue that, in conventional CaPO₄/collagen biocomposites, there is no chemical bonding between phases, possibly due to inadequate interfacial bonding during mixing [766]. Nevertheless, this is not the case with phosphorylated collagen [749]. Experts in computer modeling refer to density functional theory studies of HA surfaces and collagen peptides [1186]. They also look to molecular dynamics simulations of HA surface interactions with three polymers (PE, PA, and PLA) [1187] and an additional three polymers (polyvinylpyrrolidone, polyacrylamide, and PVA) [1188].

Researchers gathered Fourier transform infrared (FTIR) spectra of several CaPO₄-based biocomposites and collagen films in order to investigate possible phase interactions [748]. The Kramers–Kronig equation was then used to convert those spectra into absorption ones, which showed energy shifts in residues at the apatite/collagen boundary. Through a thorough comparison of the biocomposite and collagen FTIR spectra, we observed red shifts in the C–O bond absorption bands within the biocomposite spectrum. The red shifts observed were explained by a decrease in the binding energy of C–O bonds and were suggested to be a result of their interaction with Ca²⁺ ions found on the surface of apatite nanosized crystals, as demonstrated in Figure 12 [748]. Another indication of the chemical interaction between apatite and collagen was evaluated in the FTIR spectra of CDHA/collagen biocomposites, whereby the bands associated with the –COO– stretching displayed a shift from 1340 to 1337 cm⁻¹ [714,715]. According to earlier research, the carboxylate groups of the collagen macromolecule interacted chemically to generate apatite crystals on collagen [1189,1190,1191].

FTIR spectroscopy appears to be a very useful technique for examining possible chemical bonding between phases in hybrid formulations and composites based on CaPO₄ [186,269,295,301,302,405,469,597,608,619,621,623,626,630,640,676,677,680,685,715,749,802,803,821,848,875,919,1192,1193,1194,1195,1196,1197,1198,1199,1200,1201]. For example, in HA/PE biocomposites, the characteristic bands of the hydrocarbon backbone of PE at 2918, 2850, and 1472 cm⁻¹ show zero shift, while, in HA/PA, bands at 3304, 1273, and 692 cm⁻¹ derived from N-H stretching, C–N–H stretching, and N–H vibrations are observed in HA/PA, which move to 3306, 1275, and 690 cm⁻¹ in biocomposites. Moreover, the stretching (3568 cm⁻¹) and vibrational (692 cm⁻¹) modes of HA hydroxide shifted to 3570 and 690 cm⁻¹ in HA/PA biocomposite, indicating the formation of hydrogen bonds. Furthermore, the bands at 1094 and 1031 cm⁻¹ in the PO₄ mode are also shifted to 1093 and 1033 cm⁻¹ in the HA/PA biocomposite. The bands’ shift in the fingerprint region suggested that hydroxides and orthophosphate on the HA surface can interact by nucleophilic addition with the plentiful carboxyl and amino groups of PA [269]. Similar conclusions were drawn for many other CaPO₄-containing composites and hybrid formulations, namely: HA/PVA [626], CDHA/alginate [715], ACP/PPF [469], HA/maleic anhydride [302], HA/carboxylated PU [1198], HA/chitosan [821,1200,1201], and β-TCP/PLLA [405] biocomposites. It has been shown that weak chemical bonds are formed between Ca²⁺ ions on the surfaces of HA, CDHA, ACP, or β-TCP, respectively, and the slightly polarized O atoms of C=O bonds in the surrounding bio-organic compounds. The data obtained showed that the crystallization of CaPO₄ in solutions containing chitosan is largely regulated by the chemical interactions of the components. Clearly, part of the calcium was retained by chitosan and was not involved in the formation of the main mineral phase [1197]. Such chemical interactions are schematically illustrated in Figure 13 [715].

Other measuring methods can also demonstrate proof of phase-to-phase chemical interactions in CaPO₄-based composites and hybrid formulations [295,405,621,622,626,1193,1194,1195,1196,1197,1198,1199,1200,1201,1202,1203]. For example, such evidence has been observed in nanosized crystals of CDHA/alendronate by thermogravimetric analysis: the DTG plots of the crystals appear quite different from those obtained from a mechanical mixture of CDHA and calcium alendronate of similar composition [1202]. Similar DTG results were obtained for nanosized HA/PVA biocomposites [626]. In the case of nanosized HA and PA biocomposites, hydrogen bonding between the phases was detected by differential scanning calorimetry [621]. Similar results were obtained in another study [1194]. Another example is the application of dynamic mechanical analysis to investigate the softening mechanism of β-TCP/PLLA biocomposites [405]. For nanosized HA and PVAP biocomposites, X-ray diffraction and thermogravimetric analysis found some indirect evidence of chemical bonding between phases [295]. The orientation of FA crystallites and gelatin within the FA/gelatin composite spheres was shown to have a strong structural link, which suggested a significant remodeling of the macromolecular matrix in the vicinity of growing aggregates [894]. It was found that reverse gas chromatography at infinite dilution can provide some data on the thermodynamic interactions between CaPO₄ and PLLA [1203]. Furthermore, using ab initio techniques, the chemical interactions between HA and organic compounds have been clarified [1204].

X-ray photoelectron spectroscopy (XPS) revealed that the binding energies of Ca, P, and O atoms were slightly different in nanosized HA (Ca: 350.5 and 345.5; O: 530.2; P: 132.5 eV) and nanosized HA/conjac-glucomannan/chitosan biocomposite (Ca: 352.1 and 347.4; O: 531.2; P: 133.4 eV), respectively [640]. Further measurements by FTIR and X-ray diffraction showed that the nanosized HA was bound to the conjugated glucomannan and chitosan mainly by hydrogen bonds between the OH⁻ and PO₄³⁻ ions of HA and the –C=O and –NH groups of the conjugated glucomannan and chitosan copolymers, forming a stable interface between the three phases of the biocomposite. On the other hand, coordinate bonds can form between Ca²⁺ and –NH. The three phases of the biocomposite form a stable interface [640]. In a different investigation, XPS revealed hydrogen connections between the surface P–OH groups of HA and the C=O groups of PDLLA [1194]. XPS was used to quantify the Ca²⁺ ions of HA and the RCOO^– groups of collagen molecules in HA/collagen biocomposites in order to generate a covalent link [598]. Several other CaPO₄-based composites and hybrid formulations have shown comparable XPS observations [619,676,677]. Figure 14 depicts conceivable types of chemical interaction that could occur between HA crystals and bioorganic molecules in HA/chitosan-gelatin network films [1196].

The possible interaction of BCP and HPMC was investigated in IBS composites [922,923,1205]. After mixing, a reduction in the average diameter of BCP granules was noted, which affected the viscosity of the paste. During the interaction, the dissolution of the grain boundaries of β-TCP crystals and accumulation of CDHA on the HA crystal surface were observed. These two phenomena are responsible for the observed particle size changes [922,923]. However, the chemical bonding of BCP to HPMC was not found within the sensitivity of the measurement technique used [1205].

CDHA/chitosan biocomposites were prepared using the co-precipitation method [810]. Organic acids with two or more carboxyl groups that were firmly bonded to the CDHA surface via COO–Ca linkages hindered the formation of CDHA crystals. Transmission electron microscopy images revealed that, while nanosized CDHA crystals were aligned along the chitosan molecules with amino groups as nucleation sites, CDHA formed elliptical aggregates due to chemical interactions (possibly coordination bonds) between Ca on the surface and amino groups of chitosan [810]. During the setting of self-hardened (TTCP + DCPA)/polyphosphazane biocomposites, calcium cross-linked polymer carboxylates were proposed to be formed; based on the pH monitoring results, the chemical participation of the polymer in the setting process was concluded [546,547,548].

For PCL/HA biocomposites prepared by the grafting method, the presence of chemical bonding between the phases was concluded [363], but, unfortunately, strong experimental evidence was not provided. In another study, CDHA/poly(α-hydroxyester) biocomposites were prepared by a chemical route at low temperatures [335]. In that study, α-TCP was combined with PLA, PLGA, and their copolymers to prepare pre-composite structures. Solvent cast or pressed pre-composites were subjected to in situ hydrolysis of α-TCP to CDHA at 56 °C, resulting in the final biocomposite. FTIR spectroscopy showed that there was no chemical reaction between the polymer and CaPO₄ during this transition [335].

In CaPO₄-based biocomposites capable of sustaining sintering at high temperatures (useful for formulations solely consisting of inorganic components), interdiffusion of chemical elements between phases may occur. Such an effect was detected by energy-dispersive X-ray spectroscopy in the partial formation of calcium titanate in HA/TiO₂ biocomposite particles. This process was found to be advantageous in increasing the cohesion of the particles in the composite coating [1206]; similar high-temperature interactions were detected between HA and zirconia [931,956], as well as between HA and Ti [661,1008,1009,1011,1019,1022]. Namely, composites with low Ti content sintered at 1200 °C exhibited the main crystalline phases of CaTiO₃, CaO, and Ti_xP_y, while Ti₂O and residual α-Ti appeared as additional phases as the Ti content increased to 50 vol%. Therefore, the chemical interaction between HA and Ti is represented by the following non-equilibrated chemical scheme [1009]:

Ti + Ca₁₀(PO₄)₆(OH)₂ → CaTiO₃ + CaO + Ti_xP_y + (Ti₂O) + (Ca₄P₂O₉) + H₂O

(1)

Such undesirable interactions between Ti and HA will be minimized if Ti particles are coated with silica [1025].

Furthermore, for HA/Al₂O₃ biocomposites, partial HA degradation and the production of different calcium aluminates were observed following sintering at 1200–1300 °C. This is believed to be caused by the diffusion of Ca²⁺ from HA into the alumina matrix and the depletion of Ca²⁺ ions from HA, which transforms HA to β-TCP. The biocomposite mechanical strength is impacted by each of these processes [964,965,966,1206].

To conclude this section, one should admit that, for many CaPO₄-containing biocomposites, the presence of some chemical bonding among the phases was found. Nevertheless, there are formulations, such as biocomposites with PE and PP, in which no chemical bonding was detected.

7. Ways to Create a Chemical Bonding among the Constituents of Biocomposites

The interaction and adhesion at the interface among the constituents have a significant impact on the properties of the composites and are essential for the transmission of loads among the phases because poor adhesion reduces the mechanical properties [301]. Therefore, an interface among the phases appears to be a very important issue in any type of composite and hybrid material. Due to this reason, sometimes, an interface is called “one of the three elements of composite material” ([1207], p. 5). When two or more materials form a composite, there are certain physical and chemical interactions among them at the interface. If they do not have compatibility to form a composite material interface, there is no bonding among them. Possible interconnections between nanosized particles and polymer chains have been classified by Kickelbick into four types (Figure 15): (1) inorganic particles embedded in inorganic polymer, (2) particle association by bonding to the polymer backbone, (3) interpenetrating networks by chemical bonding, and (4) inorganic–organic hybrid polymers [1208].

Nevertheless, the interactions between phases are mechanical in character in a large portion of the above-mentioned formulations. This is due to the fact that the matrix is typically made up of substances devoid of unsaturated bonds or functional groups that could allow CaPO₄ components to form ionic complexes. It is evident that the non-polar polymer and the CaPO₄ ceramic particles have little in common. Thus, more promising in this regard are polymers having functional groups suspended on the polymer backbone that serve as bridges to CaPO₄ [337].

Various approaches have been proposed to solve this problem. Namely, researchers pressed blends with varying PLLA and HA concentrations at varying temperatures and pressures in order to optimize the filler’s adherence to the matrix [1209]. About 15 weight percent of PLLA provided the maximum compressive strength; however, by using blends containing 20 wt% PLLA, the authors observed that the mechanical properties were improved with increasing pressing temperature and pressure. The first finding was attributed to the increased wettability of HA particles due to the decrease in the viscosity of PLLA with increasing temperature. The second one was explained by the increased fluidity of the polymer at higher temperatures and increased compression and pore penetration at higher pressures. The combination of high pressure and high temperature has been found to reduce porosity, guarantee particle-polymer adhesion, and increase the compressive strength [277] and fracture energy [1210] of CaPO₄-containing biocomposites. Interestingly, to improve an interfacial bonding between PLLA and HA by stereocomplexation, PDLA was initially grafted onto the surface of HA particles, and then the PDLA-grafted HA particles were incorporated into a PLLA matrix [1211]. In addition, various polymer crosslinking methods have been established to tailor the chemistry of the organic and CaPO₄ components and keep the mechanical properties of the biocomposites constant [401].

However, using auxiliary reagents that can positively influence the interaction between the phases appears to be the most common practice. Such additives are called coupling agents [1207,1212]. All coupling agents contain two special functional groups that can be connected with two different types of materials and promote adhesion among the phases. Frequently, the effect is non-specific and can affect, for example, the rheology of the composite [268]. For example, diisocyanate coupling agents [282,1213] and other polymers [1192] were used to bind PEG/PBT (Polyactive^TM) block copolymers to HA filler particles. Other researchers modified a surface of nanodimensional HA with dopamine and hexamethylenediamine and grafted PLLA onto it by an aminolysis reaction. Afterward, the PLLA-grafted HA particles were blended with PLLA to fabricate PLLA/HA scaffolds by 3D printing. The results showed that the coating of the polydopamine layer introduced amino groups onto the surface of HA, which acted as active sites for grafting PLLA chains by aminolysis reaction and enhanced interfacial bonding between HAP and PLLA [1214]. Similar results with other coupling agents were obtained in other studies [1215,1216,1217].

Thus, surface modification of the grafted HA filler significantly increased the modulus and strength of polymers compared to polymers filled with ungrafted HA [282,1213,1214,1215,1216]. The polymers were chemically attached to the HA particles via isocyanate groups, as demonstrated by thermogravimetric and infrared studies. This is an appropriate strategy to increase adhesion [1192]. Other researchers have used glutaraldehyde as a crosslinking reagent [415,595,597,598,613,738,779,783,1218]. Silanes [258,259,282,351,625,1219,1220,1221,1222,1223,1224], zirconates [268,351,353,1225,1226,1227], titanates [268,351,1226,1228], phosphoric acid [629], alkaline pretreatment [867,868], PAA [1229], hexamethylene diisocyanate [1230], and other chemicals [1230] were found to have the ability to promote the interfacial bonding between CaPO₄ and other components. Additionally, some polymers can be grafted onto the surface of CaPO₄ particles [648,1230]. Namely, HA particles were modified by hexamethylene diisocyanate and cross-linked with PVA in situ to prepare a hydrogel suitable for a potential application as a calcified cartilage layer [1230].

Structural modification of the polymer matrix by, for example, adding acrylic acid [258,259,282], has also been found to be an effective method. In HA/Polyactive^TM composites, for instance, the use of polyacids as binding agents produced surface-modified HA particles with enhanced mechanical qualities and greater interaction with the polymer during fracture [282]. It was discovered that the molding technique had a significant impact on the employment of titanate and zirconate binding agents [268]. Prior to being used as fillers in biodegradable composites, silane-bonded HA powders underwent testing. This process allowed HA to withstand aqueous attack without compromising its overall biological activity [1220,1221,1222,1223,1224]. Furthermore, chemically modified reinforced phase–matrix interfaces were found to improve the mechanical properties of biocomposites. Examples include chemically bonded HA/PE [258,259], chemically formed HA/Ca poly(vinyl phosphonate) [299], and PLA/HA fibers [235]. Those biocomposites can consume large amounts of energy in the fracture.

The effect of some binding agents was found to combine two different mechanisms: (i) cross-linking of the polymer matrix (effective with zirconate and titanate binding agents) and (ii) enhanced interfacial interaction between the main phases of the biocomposite. This enhanced interfacial adhesion was found to be highly dependent on the chemistry of the binding agent (pH and type of metal center) [351]. Several studies have claimed that silanes interact with HA [258,289,1220,1221,1222], suggesting the presence of a silicon-containing interfacial phase between HA and PE that facilitates chemical adhesion between HA particles and polymers. Additionally, silane coupling agents encouraged PE to enter the spaces between individual HA particles, which strengthened the mechanical interlocking at the matrix–supporter contact [258,259].

Therefore, optimizing the properties of biocomposites with binding agents is now an important area of research. It has been suggested that controlling and enhancing the coupling between polymers and CaPO₄ at the molecular level is important for the resulting mechanical response of biocomposites. The fundamental molecular understanding of interfacial behavior in biocomposites appears to be an under-addressed area in the literature. Various experimental characterization techniques using electron microscopy, vibrational spectroscopy, X-ray diffraction, and scanning probe microscopy are routinely used to characterize such materials beyond mechanical characterization. In addition, atomic-scale models to simulate phase interactions and predict responses in novel material systems, including nanostructures and nanointerfaces, are important for understanding and predicting load-deformation behavior [1172].

Moreover, the CaPO₄ surface itself can be modified in a similar way [648,1225,1226,1227,1228,1229,1230,1231,1232,1233,1234,1235,1236,1237,1238,1239,1240]. This process is also known as CaPO₄ surface functionalization [1238]. A fascinating method for modifying the HA surface has been reported [1236]. First, during the hydrothermal synthesis of HA, 3-mercaptopropionic acid was used to achieve in situ synthesis of surface thiol-functionalized HA (HA-SH) (A in Figure 16). This was followed by the generation of sulfur-centered radicals on the HA surface by radical chain transfer (B in Figure 16) initiating surface graft polymerization of ethylene glycol methacrylate phosphate (C in Figure 16) [1236]. In one case, surface functionalization of CaPO₄ particles was found to reduce bacterial adhesion to their surface [1237]. Other examples can be found in the literature [648,1225,1226,1227,1228,1229,1230,1231,1232,1233,1234,1235,1236,1237,1238,1239,1240]. In general, the aim of surface modification is to prevent or delay the exfoliation process of CaPO₄ particles from the matrix and at the same time ensure homogeneous distribution of CaPO₄ particles in the matrix at high loading levels. Naturally, all surface modifiers must meet various biomedical requirements, such as non-toxicity, good biocompatibility, and no alteration of the biological or physicochemical properties of the filler [1238].

The addition of adhesion promoters may become an alternative way to improve the interaction between the filler and matrix. For example, 4-methacryloyloxyethyl trimellitate anhydride was added to promote polymer adhesion to HA [1241]. In another study, phosphate esters were added to the liquid component of the formulation [1242]. Both the strength and the affinity index of the biocomposite were found to increase, probably due to copolymerization.

In almost all studies on HA/carbon nanotube biocomposites, the nanotubes have been functionalized before incorporation into HA. Most researchers have achieved this via oxidation [180,181,182,1093,1094], but non-covalent functionalization with sodium dodecyl sulfate [1093] and coating the nanotubes with a polymer prior to incorporation into HA [1243] have also been reported. Several transmission electron microscopy studies have shown that functionalization enhances the interaction between carbon nanotubes and HA [1093,1094,1244].

To conclude, there are several ways to either improve or create chemical bonding among the constituents of CaPO₄-containing composites and hybrid formulations, but using coupling agents currently seems to be the most common way to solve the problem. However, the reactive sites on the CaPO₄ surface are so many that the content of the coupling agent is very difficult to control. According to the topical publication: “Too much coupling agent may condense with adjacent silanols to form a siloxane layer or remain partly uncondensed at the surface, which thereby decreasing the mechanical properties. Therefore, the content of the coupling agent should be controlled carefully to maximize the efficiency of interfacial reinforcement. Additionally, the coupling agent is a chemical agent, which must be studied widely in vitro and in vivo for its biocompatibility before being used in the human body” ([1207], p. 13).

8. Bioactivity and Biodegradability of CaPO₄-Containing Biocomposites

All types of biodegradable formulations begin to decompose and resorb in vivo by surrounding fluids and cells after implantation. The biodegradation process is accompanied by loss of mass and changes in shape and size. In general, biodegradation of implants occurs in two ways: surface erosion and bulk degradation. Surface erosion occurs mainly in the outermost layer and at the interface between the implant and the surrounding tissue, while bulk degradation occurs simultaneously in the entire volume of the implant. Therefore, biodegradation appears to be a combination of these two processes with different weights, but the underlying mechanisms depend on the type of material [1245].

The bioactivity and biodegradability of composites and hybrid formulations are therefore determined by the same properties of the components. These two processes are highly multifactorial, as the surface of the graft encounters biological fluids after implantation and is colonized by cells shortly afterward. This highly complex process involves much more biology than chemistry and materials science [1246], and many specific details are not yet known. In addition, biodegradation of all components of a biocomposite can occur simultaneously and the resulting by-products can affect both the overall process and the biodegradation of individual components. For example, in biocomposites prepared from polyesters and TCP, hydrolysis reactions of ester bonds, acid decomposition of carboxyl end groups, dissolution of TCP, and buffering reactions due to dissolved phosphate ions occur simultaneously [1247,1248,1249,1250,1251]. Swelling was also detected [1246]. In those formulations, autocatalysis is inhibited, and polymer degradation is postponed by basic TCP buffering the polyester’s acidic breakdown products. As a result, the pure polymer sample experiences a reduction in pH and mass at an earlier stage of degradation than the matching composite. This is not always the case, though. Namely, there have been studies showing that the presence of CaPO₄ does not affect the degradation rate of polymer matrices [1252,1253,1254]. Therefore, to simplify the learning, the biodegradation of individual components needs to be considered independently. In other words, the chemical dissolution of physiologically significant CaPO₄ in weakly acidic media (they are nearly insoluble in alkaline solutions [74,75]) can account for the in vitro biodegradation of the compound. In the case of CDHA, this can be expressed as four successive chemical Equations (2)–(5) [1255,1256]:

Ca_10−x(HPO₄)_x(PO₄)_6−x(OH)_2−x + (2 − x)H⁺ = Ca_10−x(HPO₄)_x(PO₄)_6−x(H₂O)_2−x^(2−x)+

(2)

Ca_10−x(HPO₄)_x(PO₄)_6−x(H₂O)_2−x^(2−x)+ = 3Ca₃(PO₄)₂ + (1 − x)Ca²⁺ + (2 − x)H₂O

(3)

Ca₃(PO₄)₂ + 2H⁺ = Ca²⁺ + 2CaHPO₄

(4)

CaHPO₄ + H⁺ = Ca²⁺ + H₂PO₄⁻

(5)

The degradation process of polymers is a chain scission process in which oligomers, monomers, and/or other low molecular weight species are formed by breaking atomic bonds, while the cleavage of polymeric bonds can be hydrolytic, enzymatic, or stimuli-associated. Therefore, the biodegradability of polymers is determined by the presence of cleavable bonds that can be broken by hydrolysis, oxidation, cellular activities, and/or enzymes, as well as under the influence of various stimuli [1245,1257]. All these types of biodegradation mechanisms of polymers are described in detail in a recent review [1258], to which the interested readers are referred. Furthermore, the biodegradability of polymers is contingent upon the subsequent factors: (1) the hydrophobicity of the monomer; (2) the chemical stability of the polymer backbone; (3) polymer morphology; (4) initial molecular weight; (5) manufacturing procedures; (6) implant geometry; and (7) porosity and pore diameter [1245,1257,1258]. The literature has a synopsis of the degradation of PLA, PGA, and SEVA-C (Ref. [125], pp. 798 and 803, in that order).

The biodegradation of HA/PLLA and CDHA/PLLA biocomposite rods in rabbit subcutaneous and medullary cavities was investigated using mechanical and histological techniques in in vivo investigations. It was found that degradation was faster when unsintered CDHA was used instead of sintered HA [1259]. Two weeks following implantation, neoplastic bone growth was seen in a more thorough investigation, particularly in preparations with a higher concentration of HA [1260]. Notably, in that instance, direct contact between these composites and bone was found, with no fiber tissue in between [1260,1261]. After 12 weeks of subcutaneous implantation, an in vivo investigation on the biodegradation of PLGA, gelatin, and PTMC microspheres/CaPO₄ biocomposites revealed that the microspheres broke down and their compressive strength decreased [1262]. Interestingly, the amount of CaPO₄ in the biocomposite was found to have a greater effect on the early stages of osteoblast behavior (cell attachment and proliferation) rather than the immediate and late stages (proliferation and differentiation) [1263]. A degradation model was also proposed for SEVA-C/HA biocomposites. It consists of three successive periods of time and, in fact, is determined by the leaching of plasticizers and low molecular weight polymeric chains with a minor HA dissolution [353].

The biodegradation of nanosized HA/collagen/PLA biocomposites was studied in 1% trypsin/phosphate-buffered saline at 37 °C for in vitro testing and the samples were implanted in the posterolateral lumbar spine of rabbits for in vivo testing [605]. The findings indicated that weight loss increased steadily in vitro, reaching a mass loss of approximately 20% after 4 weeks. The relative rates of reduction of the three components in this material varied significantly during the experimental period in vitro: collagen decreased the fastest, going from 40% by weight to about 20% in the composite; the HA content increased from 45 to approximately 60%; and the amount of PLA changed slightly. It was discovered that, in vivo, the collagen/HA ratio was marginally greater closer to the transverse processes than it was in the space between them [605]. High-strength HA/PLLA biocomposite rods were used for internal fixation of fractures in a 5–7-year in vivo study for internal fixation of fractures [1264]. In that study, both unsintered CDHA and sintered HA were used as reinforcing phases within a PLLA matrix. The prepared biocomposites were implanted into the femurs of 25 rabbits. The results showed that the implanted material resorbed six years after implantation. Bone remodeling and the presence of trabecular bone union were significant results [1264]. Data on the release kinetics of monomers from dental composites containing fluoride-doped CaPO₄-containing biocomposites are also available [1265].

To conclude this part, one should mention an outstanding Ph.D. thesis defended by Dr. Ismael Moreno-Gomez in 2019 at Cambridge University devoted to the mathematical modeling of the degradation of bioresorbable composites [1266], in which chapters 4 and 5 are devoted to the degradation of the TCP- and HA-containing biocomposites, respectively.

9. Biomedical and Non-Biomedical Applications

First, one should remember that the goal of this review is to inform readers of a wide variety of known CaPO₄-containing composites and hybrid formulations. Even cataloging them and providing a brief description of their composition and properties coupled with the key processing parameters appears to be rather long. Therefore, let me limit myself to just brief descriptions of the most important applications.

Due to the fact that CaPO₄ represents the major inorganic constituent of calcified tissues of mammals, most CaPO₄-based composites and hybrid materials have been developed for potential applications as artificial bone grafts. That is why they are called biocomposites and biomaterials. In addition to all that has been previously mentioned, the examples of biomedical applications of CaPO₄-containing formulations comprise regeneration of critical-sized mandibular defects [1267], orthopedic applications [401], bone revascularization [1268], bone tissue engineering [1269,1270], anti-cancer treatments [1064,1065], etc. Additionally, composite fillers composed of HA and CMC gel possess biostimulatory and skin-tightening properties that have been applied not just to the face but also to the body to boost rejuvenation [900,901].

Non-biomedical applications of CaPO₄-based composites and hybrid formulations are also known, and their amount rapidly increases [460,461,834,835,836,874,980,1089,1090,1271,1272,1273,1274,1275,1276,1277,1278,1279,1280,1281,1282,1283,1284,1285]. For example, battery devices employ DCPD/polymer composites as proton conductors [460,461], an HA-contained porous gel polymer electrolyte was developed for a quasi-solid-state sodium ion battery [1280], and HA/carbon nanotube was found to be an anode modifying material [1284], while CuO/HA composites appeared to be excellent dielectric materials [1283]. Other formulations composed from CaPO₄ and wood processing residues, agricultural fibers (such as dried crushed sugarcane bagasse), pulp mill sludge, or waste-paper were found to meet the requirements for Portland-cement-bonded particleboards [1271]. In addition, nano-fibrillated cellulose/HA composites appeared to possess excellent fire [1273] and water [874] resistances, while both HA/calcium silicate [1274] and HA/manganese dioxide [980] were found to be effective adsorbents for Pb removal from aqueous solutions. All-weather flexible electrically conductive paper [1272] and another recyclable, fire-resistant, superhydrophobic, and magnetic paper [1278] were produced using nanodimensional HA wires, while a heat-insulating and fire-retardant inorganic paper was produced by employing ultralong HA nanowires as scaffolds for silica aerogels particles [1277]. HA/gold composites appear to preserve some wooden artifacts [1276]. Furthermore, PEG/HA composites appeared to be a suitable phase change material, which was useful for thermal energy storage [1275], while a full-daytime sub-ambient cooling TCP/acrylic paint was found to be effective for energy saving [1281]. In addition, CaPO₄/chitosan composites could be used as a plant growth promoter [1282], while HA/gelatin ones appeared to resemble a synthetic ivory suitable to produce piano keys [1279]. Finally, there are different fields of application for bone char, starting from environmental remediation to sustainable agriculture and catalysis [1089,1090].

However, the most unusual application of CaPO₄-based composites and hybrid formulations lies at the interface between biomedical and non-biomedical applications. Namely, there is a recent study in which simple, translucent, highly stretchable, and high-performance HA/polydimethylsiloxane bionanocomposite film-based triboelectric nanogenerators were developed to synergistically capture external mechanical energy from biomechanical sources, such as wearable devices, into electricity [1286]. A triboelectric nanogenerator (2 × 2 cm²) with a pushing force of 2 N and different amounts of HA in polydimethylsiloxane produced a highly stable output voltage, current, surface charge density, and power density values of 300 V, 22.4 μA, 90.36 μC/m², and 27.34 W/m², which were 6-, 9-, and 10-times higher than those without HA, respectively. The HA/polydimethylsiloxane bionanocomposites were found to exhibit a remarkable stretchability of more than 290%. The authors attached these nanogenerators to cloth, skin, an insole, and a person’s lap. Effectively harvesting energy from body movements, the triboelectric nanogenerators may be used to charge multiple commercial capacitors, drive up to 100 LEDs, and power a low-power electronic device, such as a commercial calculator. Therefore, self-powered sensing and wearable devices are made possible by HA/polydimethylsiloxane biocomposite nanogenerators, which allow their large-scale preparation and deployment [1286].

10. Some Issues and Key Challenges

The objective of this review is to summarize scientific information on various fabrication approaches that have been used to produce a broad range of CaPO₄-based composites and hybrid formulations. These materials aim to enhance the weak properties of CaPO₄ bioceramics, specifically mechanical ones, without compromising their biological performances. However, reasonable compromises are often necessary. Nevertheless, these biocomposites, particularly if enhanced by biologically active biocompounds like osteoconductive and osteoinductive factors, as well as osteogenic cells, are already of significant interest as a novel group of adaptable biomaterials and are acknowledged for their usefulness in numerous bone grafting applications [22,1287]. Regrettably, despite the significant development and testing of CaPO₄-based composites and hybrid formulations in numerous biomedical applications, their commercial distribution and industrial production fall short of expectations. Presumably, this stems from the fact that each formulation and fabrication technique has specific advantages and disadvantages. As a result, there are still disputes that need to be explored further in bone tissue engineering to obtain suitable CaPO₄-based biocomposite scaffolds with tunable properties. Therefore, better comprehension of their performance and limitations is necessary for their biomedical applications, which require further in vivo investigations. The main issues are summarized as follows [251]:

• When compared to traditional monolithic materials, the long-term performance of biocomposites is not well-supported by sufficient credible experimental and clinical data.
• The number of design variables that must be taken into account makes the design of hybrid and composite formulations significantly more challenging than the design of classic single-phase ones.
• Manufacturing techniques might be costly, time-consuming, dependent on cutting-edge technology, and restricted to certain reinforcing configurations. They might also call for certain cleaning and sterilizing procedures.
• There are currently no adequate guidelines for biocompatibility testing of biocomposite implants since it is unclear how the components of biocomposites interact with biological tissues.
• The fatigue behavior of biocomposites is much more complex and less predictable than conventional single-phase materials and there are not enough criteria to evaluate the fatigue performance of the biocomposites.

Significant research remains to be conducted regarding the analysis of cells and their various behaviors in relation to interacting with CaPO₄-containing biocomposites. Crucial unresolved questions include the following: (1) What mechanisms underlie the promotion of cellular growth and differentiation by these biocomposites? (2) How can these pathways be determined? In the future, biocomposite surfaces will be functionalized with molecules of various properties and sizes that can selectively interact with biomolecules, such as proteins and peptides, by binding to cells [218].

Although processing techniques for biocomposites have advanced significantly, there is still a need for more sophisticated technological advances to produce bone-like hierarchical structures at various length scales to achieve the desired properties. The development of new implantable materials relies on advances in research on natural bone structure. One of the primary obstacles is to comprehend the basis of biomineralization and then effectively apply this knowledge to develop superior synthetic methods for producing bone graft materials. Unfortunately, the production of biocomposites that mimic natural bone at nanometer to micrometer levels presents numerous poorly understood issues, such as morphological regulation, the addition of foreign ions, interactions with proteins, and the construction of organic and inorganic phases. The gap in processing between lower-grade constituent units and higher-grade architectures may considerably restrict the practical applications of existing CaPO₄-based composites and hybrid formulations. Thus, important additional research endeavors are outlined to tackle the ensuing crucial challenges [22]:

• Optimization of processing conditions for biocomposites.
• Enhancement of interfacial adhesion and strength similar to that of genuine bone.
• To facilitate bone formation, both surface characteristics and pore size should be optimized.
• Maintaining sufficient in vivo structure volume for osteogenesis to occur.
• Resistance to load-bearing conditions.
• Matching the bioabsorbability and biomechanical properties of the graft when creating new bone.
• Understanding the molecular processes by which cells and biocomposite matrices interact in vivo to support bone regeneration is necessary.
• Promoting angiogenesis and vascularization for healthy bone cell growth and subsequent tissue formation and remodeling.

To widely commercialize CaPO₄-based composites and hybrid formulations in surgery and medicine, these important difficulties must be resolved.

11. Conclusions

The calcified tissues in mammals contain poorly crystalline ion-substituted CaPO₄ with an apatitic structure (CDHA) and have a complex hierarchical assembly, making them complicated bioorganic/inorganic composites. Their mechanical properties are remarkable, considering the weak components they consist of, surpassing what can be achieved by currently available synthetic materials. This is due to the fact that organisms produce biocomposites that are hierarchically ordered at the nano, micro, and meso levels and comprise brittle CaPO₄ and ductile bio-organic components. These biocomposites are organized in both composition and structure. Furthermore, calcified tissues always have multiple functions, therefore necessitating multifunctional CaPO₄-containing composites and hybrid formulations [1288]. Bone provides structural support for the body and is responsible for hematopoiesis. Additionally, in contrast to current synthetic systems, biological systems possess the capacity for self-repair, which is almost universal in nature. These intricate structures, developed over millions of years of evolution, have motivated materials scientists to create novel biomaterials [1289].

The present review highlights a wide range of CaPO₄-containing composites and hybrid formulations currently available for biomedical applications. By synergizing the favorable biocompatible properties of CaPO₄ with the mechanical ones of other compounds, a significant potential arises for the advancement of bone substitute materials with markedly improved microstructure and properties. Therefore, it is essential to develop multiphase biocomposites that offer a multistage design approach and consequently offer more control over their material and biological properties than the individual constituents alone. As a result, the addition of CaPO₄ to polymers was conceived, leading to the development of bioactive biodegradable biocomposites from the originally bioinert monolithic materials used in the 1970s and 1980s. These materials are now utilized as elasticity-matched biomaterials and recently as scaffolds for tissue engineering [1290]. The application of traditional composite manufacturing techniques to the biomaterials field has resulted in this approach. The review’s summary of several studies shows that there are several benefits to using a brittle, tough, bioactive CaPO₄ filler in combination with a ductile matrix for biomedical applications. The favorable characteristics of each component can offset the inadequate mechanical performance of CaPO₄ bioceramics. Furthermore, the beneficial bioactive features of CaPO₄ can improve the qualities of the other phase, broadening the in vivo usage capabilities of both materials [66,67]. However, the literature review indicates that researchers have primarily investigated simple, complex, and graded CaPO₄-based composites and hybrid formulations, as well as fibrous, laminar, and particulate ones (refer to Section 2 for the classification types of composites). CaPO₄-containing composites and hybrid formulations have already been developed in various forms, including scaffolds [1291], coatings [1292,1293], injectable [567,904], self-hardening formulations [478], and concretes [517,518]. However, in the future, further attention should be given to elaborating and developing both hierarchical and hybrid biocomposites. In addition, the research will continue to prioritize shape memory biocomposites [389,390,391], further developments of functionally graded composites [1294], biocomposites with novel properties, such as piezoelectric ones [1295], and biocomposites with therapeutic properties for bone tissue engineering applications. These innovative formulations offer multiple benefits, including reduced incision length, mimicking of bone porosity, controlled and increased strength, and potential antibacterial and anti-inflammatory effects. Furthermore, in accordance with current trends in tissue engineering, next-generation CaPO₄-based composites and hybrid formulations should incorporate various biological components, such as hormones, bioactive factors, drugs, and living cells [1269,1296,1297]. These components could augment the biomaterials’ osteoconductivity, osteointegration potential, and biocompatibility, thus making them suitable for biomedical applications.