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Dorozhkin, S.V. Calcium Orthophosphate-Based Bioceramics. Encyclopedia. Available online: https://encyclopedia.pub/entry/27794 (accessed on 07 July 2024).
Dorozhkin SV. Calcium Orthophosphate-Based Bioceramics. Encyclopedia. Available at: https://encyclopedia.pub/entry/27794. Accessed July 07, 2024.
Dorozhkin, Sergey V.. "Calcium Orthophosphate-Based Bioceramics" Encyclopedia, https://encyclopedia.pub/entry/27794 (accessed July 07, 2024).
Dorozhkin, S.V. (2022, September 28). Calcium Orthophosphate-Based Bioceramics. In Encyclopedia. https://encyclopedia.pub/entry/27794
Dorozhkin, Sergey V.. "Calcium Orthophosphate-Based Bioceramics." Encyclopedia. Web. 28 September, 2022.
Calcium Orthophosphate-Based Bioceramics
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This entry is adapted from the peer-reviewed paper 10.3390/coatings12101380

Various types of materials have been traditionally used to restore damaged bones. In the late 1960s, a strong interest was raised in studying ceramics as potential bone grafts due to their biomechanical properties. A short time later, such synthetic biomaterials were called bioceramics. During the past, there have been a number of important achievements in this field. Namely, after the initial development of bioceramics that was just tolerated in the physiological environment, an emphasis was shifted towards the formulations able to form direct chemical bonds with the adjacent bones. Afterwards, by the structural and compositional controls, it became possible to choose whether the CaPO4-based implants would remain biologically stable once incorporated into the skeletal structure or whether they would be resorbed over time. At the turn of the millennium, a new concept of regenerative bioceramics was developed, and such formulations became an integrated part of the tissue engineering approach. Now, CaPO4-based scaffolds are designed to induce bone formation and vascularization. These scaffolds are usually porous and harbor various biomolecules and/or cells.

calcium orthophosphates hydroxyapatite tricalcium phosphate

1. Bioceramics of CaPO4

1.1. History

The detailed history of HA and other types of CaPO4, including the subject of CaPO4 bioceramics, as well as description of their past biomedical applications, might be found elsewhere [1][2], where the interested readers are referred. One should just note that the earliest book devoted to CaPO4 bioceramics was published in 1983 [3].

1.2. Chemical Composition and Preparation

Currently, CaPO4 bioceramics can be prepared from various sources [4][5][6][7][8][9][10][11][12][13]. Nevertheless, up to now, all attempts to synthesize bone replacement materials for clinical applications featuring the physiological tolerance, biocompatibility, and a long-term stability have had only relative success; this clearly demonstrates both the superiority and a complexity of the natural structures [14].
In general, a characterization of CaPO4 bioceramics should be performed from various viewpoints such as the chemical composition (including stoichiometry and purity), homogeneity, phase distribution, morphology, grain sizes and shape, grain boundaries, crystallite size, crystallinity, pores, cracks, surface roughness, etc. From the chemical point of view, the vast majority of CaPO4 bioceramics are based on HA [15][16][17][18][19], both types of TCP, and various multiphasic formulations thereof [20]. Biphasic formulations (commonly abbreviated as BCP–biphasic calcium phosphate) are the simplest among the latter ones. They include β-TCP + HA [21][22][23][24][25][26][27][28][29], α-TCP + HA [30][31][32], and biphasic TCP (commonly abbreviated as BTCP), consisting of α-TCP and β-TCP [33][34][35][36][37][38]. In addition, triphasic formulations (HA + α-TCP + β-TCP) have been prepared as well [39][40][41][42]. Further details on this topic can be found in a special research [20]. Leaving aside a big subject of DCPD-forming self-setting formulations [43][44], one should note that just a few publications on bioceramics prepared from other types of CaPO4 are available.
The preparation techniques of various types of CaPO4 have been extensively research in the literature [45][46][47][48][49][50], where the interested readers are referred. Briefly, when compared to both α- and β-TCP, HA is a more stable phase under the physiological conditions, as it has a lower solubility and, thus, slower resorption kinetics [51][52][53]. Therefore, the BCP concept is determined by the optimum balance of a more stable phase of HA and a more soluble TCP. Due to a higher biodegradability of the α- or β-TCP component, the reactivity of BCP increases with the TCP/HA ratio increasing. Thus, in vivo bioresorbability of BCP can be controlled through the phase composition [22]. Similar conclusions are also valid for the biphasic TCP (in which α-TCP is a more soluble phase), as well as for both triphasic (HA, α-TCP, and β-TCP) and yet more complex formulations [20].
As implants made of sintered HA are found in bone defects for many years after implantation, bioceramics made of more soluble types of CaPO4 are preferable for the biomedical purposes. Furthermore, the experimental results showed that BCP had a higher ability to adsorb fibrinogen, insulin, or type I collagen than HA [54]. Thus, according to both observed and measured bone formation parameters, CaPO4 bioceramics have been ranked as follows: low sintering temperature BCP (rough and smooth) ≈ medium sintering temperature BCP ≈ TCP > calcined low sintering temperature HA > non-calcined low sintering temperature HA > high sintering temperature BCP (rough and smooth) > high sintering temperature HA [55]. This sequence was developed in the year 2000 and, thus, neither multiphase formulations nor other CaPO4 are included.

1.3. Forming and Shaping

In order to fabricate CaPO4 bioceramics in progressively complex shapes, scientists are investigating the use of both old and new manufacturing techniques. These techniques range from an adaptation of the age-old pottery techniques to the newest manufacturing methods for high-temperature ceramic parts for airplane engines; namely, reverse engineering [56][57] and rapid prototyping [58][59][60] technologies have revolutionized a generation of physical models, allowing the engineers to efficiently and accurately produce physical models and customized implants with high levels of geometric intricacy. Combined with computer-aided design and manufacturing (CAD/CAM), complex physical objects of the anatomical structure can be fabricated in a variety of shapes and sizes. In a typical application, an image of a bone defect in a patient can be taken and used to develop a three-dimensional (3D) CAD computer model [61][62][63][64][65]. Then, a computer can reduce the model to slices or layers. Afterwards, 3D objects and coatings are constructed layer-by-layer using rapid prototyping techniques. The examples comprise fused deposition modeling [66][67], selective laser sintering [68][69][70][71][72][73], laser cladding [74][75][76][77], 3D printing and/or plotting [78][79][80][81][82][83][84][85], robocasting [86][87][88], solid freeform fabrication [89][90][91][92][93][94], stereolithography [95][96][97][98], and direct light processing [99]. More advanced techniques, such as 4D [100][101] and 5D [102] printing techniques, have been introduced as well. Three-dimensional printing of the CaPO4-based self-setting formulations is known as well [83]. Additional details of these techniques are available in the literature [103][104][105][106].
In addition to the aforementioned modern techniques, classical forming and shaping approaches are still widely used. The selection of the desired technique depends greatly on the ultimate application of the bioceramic device, e.g., whether it is for a hard-tissue replacement or an integration of the device within the surrounding tissues. In general, three types of processing technologies might be used: (1) employment of a lubricant and a liquid binder with ceramic powders for shaping and subsequent firing; (2) application of self-setting and self-hardening properties of water-wet molded powders; (3) materials are melted to form a liquid and are shaped during cooling and solidification [107][108][109]. Since CaPO4 are either thermally unstable (MCPM, MCPA, DCPA, DCPD, OCP, ACP, CDHA) or have a melting point at temperatures exceeding ~1400 °C with a partial decomposition (α-TCP, β-TCP, HA, FA, TTCP), only the first and the second consolidation approaches are used to prepare bulk bioceramics and scaffolds. The methods include uniaxial compaction [86][110][111], isostatic pressing (cold or hot) [28][112][113][114][115][116][117][118], granulation [119][120][121][122][123][124][125], loose packing [126], slip casting [127][128][129][130][131][132][133], gel casting [95][134][135][136][137][138][139], pressure mold forming [140][141][142], injection molding [143][144][145][146], polymer replication [147][148][149][150][151][152][153][154], ultrasonic machining [155], extrusion [156][157][158][159][160][161][162], and slurry dipping and spraying [163]. In addition, to form ceramic sheets from slurries, tape casting [135][164][165][166][167][168], doctor blade [169], and colander methods can be employed [107][108][109]. In addition, flexible, ultrathin (of 1 to several microns thick), freestanding HA sheets were produced by a pulsed laser deposition technique, followed by thin film isolation technology [170]. Various combinations of several techniques are also possible [135][171][172][173][174]. Furthermore, some of those processes might be performed under the electromagnetic field, which helps crystal aligning [129][132][175][176][177][178]. Finally, the prepared CaPO4 bioceramics might be subjected to additional treatments (e.g., chemical, thermal, and/or hydrothermal ones) to convert one type of CaPO4 into another one [154].
To prepare bulk bioceramics, powders are usually pressed damp in metal dies or dry in lubricated dies at pressures high enough to form sufficiently strong structures to hold together until they are sintered [179]. An organic binder, such as polyvinyl alcohol, helps to bind the powder particles altogether. Afterwards, the binder is removed by heating in air to oxidize the organic phases to carbon dioxide and water. Since many binders contain water, drying at ~100 °C is a critical step in preparing damp-formed pieces for firing. Too much or too little water in the compacts can lead to blowing apart the ware on heating or crumbling, respectively [107][108][109][113]. Furthermore, removal of water during drying often results in subsequent shrinkage of the product. In addition, due to local variations in water content, warping and even cracks may be developed during drying. Dry pressing and hydrostatic molding can minimize these problems [109]. Finally, the manufactured green samples are sintered.
It is important to note that forming and shaping of any ceramic products require a proper selection of the raw materials in terms of particle sizes and size distribution; namely, tough and strong bioceramics consist of pure, fine, and homogeneous microstructures. To attain this, pure powders with small average size and high surface area must be used as the starting sources. However, for maximum packing and least shrinkage after firing, mixing of ~70% coarse and ~30% fine powders have been suggested [109]. Mixing is usually carried out in a ball mill for uniformity of properties and reaction during subsequent firing. Mechanical die forming or sometimes extrusion through a die orifice can be used to produce a fixed cross-section.
Finally, to produce the accurate shaping, necessary for the fine design of bioceramics, machine finishing might be essential [63][107][180][181]. Unfortunately, cutting tools developed for metals are usually useless for bioceramics due to their fragility; therefore, grinding and polishing appear to be the most convenient finishing techniques [63][107]. In addition, the surface of CaPO4 bioceramics might be modified by various supplementary treatments [182][183], and CaPO4 bioceramics might be subjected to post-processing actions, such as immersing into special solutions [184].

1.4. Sintering and Firing

After being formed and shaped, the CaPO4 bioceramics are commonly sintered. A sintering (or firing) procedure is a thermal process in which loosely bound particles are converted into a consistent solid mass under the influence of heat and/or pressure without melting the particles. This process is of great importance to manufacture bulk bioceramics with the required mechanical properties. Usually, this technique is carried out according to controlled temperature programs of electric furnaces in adjusted ambience of air with necessary additional gasses; however, always at temperatures below the melting points of the materials. The firing step can include temporary holds at intermediate temperatures to burn out organic binders [107][108][109]. The heating rate, sintering temperature, and holding time depend on the starting materials. For example, in the case of HA, these values are in the ranges of 0.5–3 °C/min, 1000–1250 °C, and 2–5 h, respectively [185]. In the majority of cases, sintering allows a structure to retain its shape. However, this process might be accompanied by a considerable degree of shrinkage [186][187][188], which must be accommodated in the fabrication process. For instance, in the case of FA sintering, a linear shrinkage was found to occur at ~715 °C and the material reached its final density at ~890 °C. Above this value, grain growth became important and induced an intra-granular porosity, which was responsible for density decrease. At ~1180 °C, a liquid phase was formed due to formation of a binary eutectic between FA and fluorite contained in the powder as impurity. This liquid phase further promoted the coarsening process and induced formation of large pores at high temperatures [189].
In general, sintering occurs only when the driving force is sufficiently high, while the latter relates to the decrease in surface and interfacial energies of the system by matter (molecules, atoms, or ions) transport, which can proceed by solid, liquid, or gaseous phase diffusion. Namely, when solids are heated to high temperatures, their constituents are driven to move to fill up pores and open channels between the grains of powders, as well as to compensate for the surface energy differences among their convex and concave surfaces (matter moves from convex to concave). At the initial stages, bottlenecks are formed and grow among the particles. Existing vacancies tend to flow away from the surfaces of sharply curved necks; this is an equivalent of a material flow towards the necks, which grow as the voids shrink. Small contact areas among the particles expand and, at the same time, a density of the compact increases and the total void volume decreases. As the pores and open channels are closed during a heat treatment, the particles become tightly bonded together, and density, strength, and fatigue resistance of the sintered object improve greatly. Grain boundary diffusion was identified as the dominant mechanism for densification [190]. Furthermore, strong chemical bonds are formed among the particles, and loosely compacted green bodies are hardened to denser materials [107][108][109]. Further knowledge on the ceramic sintering process can be found elsewhere [191].
In the case of CaPO4, the earliest research on their sintering was published in 1971 [192]. Since then, numerous research on this subject have been published, and several specific processes have been found to occur during CaPO4 sintering. Firstly, moisture, carbonates and all other volatile chemicals remaining from the synthesis stage, such as ammonia, nitrates, and any organic compounds, are removed as gaseous products. Secondly, unless powders are sintered, the removal of these gases facilitates production of denser ceramics with subsequent shrinkage of the samples. Thirdly, all chemical changes are accompanied by a concurrent increase in crystal size and a decrease in the specific surface area. Fourthly, a chemical decomposition of all acidic orthophosphates and their transformation into other phosphates (e.g., 2HPO42− → P2O74− + H2O) takes place. In addition, sintering causes toughening [18], densification [19][193], partial dehydroxylation (in the case of HA) [19], a partial evaporation and condensation of phosphates [194], and grain growth [190][195], as well as a mechanical strength increasing [196][197][198]. The latter events are due to presence of air and other gases filling gaps among the particles of unsintered powders. At sintering, the gases move towards the outside of powders, and green bodies shrink owing to decrease of distances among the particles. For example, sintering of biologically formed apatites was investigated [199][200] and the obtained products were characterized [201][202]. In all cases, the numerical value of the Ca/P ratio in sintered apatites of biological origin was higher than that of the stoichiometric HA. One should mention that in the vast majority of cases, CaPO4 with Ca/P ratio < 1.5 are not sintered, since these compounds are thermally unstable, while sintering of nonstoichiometric CaPO4 (CDHA and ACP) always leads to their transformation into various types of biphasic, triphasic, and multiphase formulations [20].
An extensive study on the effects of sintering temperature and time on the properties of HA bioceramics revealed a correlation between these parameters and density, porosity, grain size, chemical composition, and strength of the scaffolds [203]. Namely, sintering below ~1000 °C was found to result in initial particle coalescence, with little or no densification and a significant loss of the surface area and porosity. The degree of densification appeared to depend on the sintering temperature, whereas the degree of ionic diffusion was governed by the period of sintering [203]. To enhance sinterability of CaPO4, a variety of sintering additives might be added [204][205][206][207].
Solid-state pressureless sintering is the simplest procedure. For example, HA bioceramics can be pressurelessly sintered up to the theoretical density at 1000–1200 °C. Processing at even higher temperatures usually lead to exaggerated grain growth and decomposition because HA becomes unstable at temperatures exceeding ~1300 °C [45][46][47][48][49][50][208][209][210]. The decomposition temperature of HA bioceramics is a function of the partial pressure of water vapor. Moreover, processing under vacuum leads to an earlier decomposition of HA, while processing under high partial pressure of water prevents the decomposition. On the other hand, the presence of water in the sintering atmosphere was reported to inhibit densification of HA and accelerate grain growth [211]. Unexpectedly, an application of a magnetic field during sintering was found to influence the growth of HA grains [195]. A definite correlation between hardness, density, and a grain size in sintered HA bioceramics was found; despite exhibiting high bulk density, hardness started to decrease at a certain critical grain size limit [212][213][214].
Since grain growth occurs mainly during the final stage of sintering, to avoid this, a new method called ‘‘two-step sintering’’ (TSS) was proposed [215]. The method consists of suppressing grain boundary migration responsible for grain growth, while keeping grain boundary diffusion that promotes densification. The TSS approach was successfully applied to CaPO4 bioceramics [27][171][216][217][218][219]. For example, HA compacts prepared from nanodimensional powders were two-step sintered. The average grain size of near full dense (>98%) HA bioceramics made via conventional sintering was found to be ~1.7 μm, while that for TSS HA bioceramics was ~190 nm (i.e., ~9 times less) with simultaneous increasing of the fracture toughness of samples from 0.98 ± 0.12 to 1.92 ± 0.20 MPa m1/2. In addition, due to the lower second-step sintering temperature, no HA phase decomposition was detected in the TSS method [216].
Hot pressing, hot isostatic pressing [28][112][117][118], or hot pressing with post-sintering [220][221], as well as “cold sintering” (which is very similar to hot pressing) [222] processes make it possible to decrease the temperature of the densification process, diminish the grain size, and achieve higher densities. This leads to finer microstructures, higher thermal stability, and subsequently better mechanical properties of CaPO4 bioceramics. In addition, microwave [223][224][225][226][227][228], spark plasma, flash [229][230], and ultrafast high-temperature [231] sintering techniques are alternative methods to the conventional sintering, hot pressing, and hot isostatic pressing. Both alternative methods were found to be time- and energy-efficient densification techniques. Further developments are still possible. For example, a hydrothermal hot pressing method was developed to fabricate OCP [232], CDHA [233], HA/β-TCP [234], and HA [220][235][236][237][238] bioceramics with neither thermal dehydration nor thermal decomposition. Further details on the sintering and firing processes of CaPO4 bioceramics are available in the literature [48][239][240].
To conclude this, one should note that the sintering stage is not always necessary. For example, CaPO4-based bulk bioceramics with the reasonable mechanical properties might be prepared by means of self-setting (self-hardening) formulations (see Section 6.1. Self-setting (Self-hardening) Formulations below). Furthermore, the reader’s attention is directed to an excellent research on various ceramic manufacturing techniques [241], in which various ceramic processing techniques are well described.

2. The Major Properties

2.1. Mechanical Properties

The modern generation of biomedical materials should stimulate the body’s own self-repairing abilities [242]. Therefore, during healing, a mature bone should replace the modern grafts and this process must occur without transient loss of the mechanical support. Unluckily for material scientists, a human body provides one of the most inhospitable environments for the implanted biomaterials. It is warm, wet, and both chemically and biologically active. For example, a diversity of body fluids in various tissues might have a solution pH varying from 1 to 9. In addition, a body is capable of generating quite massive force concentrations, and the variance in such characteristics among individuals might be enormous. Typically, bones are subjected to ~4 MPa loads, whereas tendons and ligaments experience peak stresses in the range of 40–80 MPa. The hip joints are subjected to an average load of up to three times the body weight (3000 N), and peak loads experienced during jumping can be as high as 10 times the body weight. These stresses are repetitive and fluctuating depending on the nature of the activities, which can include standing, sitting, jogging, stretching, and climbing. Therefore, all types of implants must sustain attacks of a great variety of aggressive conditions [243]. Regrettably, there is presently no artificial material fulfilling all these requirements.
Now it is important to mention that the mechanical behavior of any ceramics is rather specific; namely, ceramics is brittle, which is attributed to high-strength ionic bonds. Thus, it is not possible for plastic deformation to happen prior to failure, as a slip cannot occur. Therefore, ceramics fail in a dramatic manner. Namely, if a crack is initiated, its progress will not be hindered by the deformation of material ahead of the crack, as would be the case in a ductile material (e.g., a metal). In ceramics, the crack will continue to propagate, rapidly resulting in a catastrophic breakdown. In addition, the mechanical data typically have a considerable amount of scatter [108]. Alas, all of these are applicable to CaPO4 bioceramics.
For dense bioceramics, the strength is a function of the grain sizes. Namely, finer-grain-size bioceramics have smaller flaws at the grain boundaries and thus are stronger than ones with larger grain sizes. Thus, in general, the strength for ceramics is proportional to the inverse square root of the grain sizes [244]. In addition, the mechanical properties decrease significantly with increasing content of an amorphous phase, microporosity, and grain sizes, while a high crystallinity, a low porosity, and small grain sizes tend to give a higher stiffness, a higher compressive and tensile strength, and a greater fracture toughness. Furthermore, ceramics strength appears to be very sensitive to slow crack growth [245]. Accordingly, from the mechanical point of view, CaPO4 bioceramics appear to be brittle polycrystalline materials for which the mechanical properties are governed by crystallinity, grain size, grain boundaries, porosity, and composition [246]. Thus, it possesses poor mechanical properties (for instance, a low impact and fracture resistances) that do not allow CaPO4 bioceramics to be used in load-bearing areas, such as artificial teeth or bones [247][248][249][250]. For example, fracture toughness (this is a property that describes the ability of a material containing a crack to resist fracture and is one of the most important properties of any material for virtually all design applications) of HA bioceramics does not exceed the value of ~1.2 MPa·m1/2 [251] (human bone: 2–12 MPa·m1/2). It decreases exponentially with increasing porosity [252]. Generally, fracture toughness increases with grain size decreasing. However, in some materials, especially noncubic ceramics, fracture toughness reaches the maximum and rapidly drops with decreasing grain size. For example, a fracture toughness of pure hot-pressed HA with grain sizes between 0.2–1.2 µm was investigated. The researchers found two distinct trends, where fracture toughness decreased with increasing grain size above ~0.4 µm and subsequently decreased with decreasing grain size. The maximum fracture toughness measured was 1.20 ± 0.05 MPa·m1/2 at ~0.4 µm [253]. Fracture energy of HA bioceramics is in the range of 2.3–20 J/m2, while the Weibull modulus (a measure of the spread or scatter in fracture strength) is low (~5–12) in wet environments, which means that HA behaves as a typical brittle ceramics and indicates a low reliability of HA implants [254]. Porosity has a great influence on the Weibull modulus [255][256]. In addition, the reliability of HA bioceramics was found to depend on deformation mode (bending or compression), along with pore size and pore size distribution: reliability was higher for smaller average pore sizes in bending but lower for smaller pore sizes in compression [257]. Interestingly, three peaks of internal friction were found at temperatures of about –40, 80, and 130 °C for HA but no internal friction peaks were obtained for FA in the measured temperature range; this effect was attributed to the differences of F− and OH− positions in FA and HA, respectively [258]. Differences in internal friction values were also found between HA and TCP [259].
Bending, compressive, and tensile strengths of dense HA bioceramics are in the ranges of 38–250, 120–900, and 38–300 MPa, respectively. Similar values for porous HA bioceramics are substantially lower: 2–11, 2–100, and ~3 MPa, respectively [254]. These wide variations in the properties are due to both structural variations (e.g., an influence of remaining microporosity, grain sizes, presence of impurities, etc.) and manufacturing processes, and they are also caused by a statistical nature of the strength distribution. Strength was found to increase with Ca/P ratio increasing, reaching the maximum value around Ca/P ~1.67 (stoichiometric HA) and decreasing suddenly when Ca/P > 1.67 [254]. Furthermore, strength decreases almost exponentially with increasing porosity [260][261]. However, by changing the pore geometry, it is possible to influence the strength of porous bioceramics. It is also worth mentioning that porous CaPO4 bioceramics are considerably less fatigue-resistant than dense ones (in materials science, fatigue is the progressive and localized structural damage that occurs when a material is subjected to cyclic loading). Both grain sizes and porosity are reported to influence the fracture path, which itself has little effect on the fracture toughness of CaPO4 bioceramics [246][262]. However, no obvious decrease in mechanical properties was found after CaPO4 bioceramics had been aged in the various solutions during the different periods of time [263].
Young’s (or elastic) modulus of dense HA bioceramics is in the range of 3–120 GPa [264][265], which is more or less similar to those of the most resistant components of the natural calcified tissues (dental enamel: ~74 GPa, dentine: ~21 GPa, compact bone: ~18–22 GPa). This value depends on porosity [266][267]. Nevertheless, dense bulk compacts of HA have mechanical resistances of the order of 100 MPa versus ~00 MPa of human bones, drastically diminishing their resistances in the case of porous bulk compacts [268]. Young’s modulus measured in bending is between 44 and 88 GPa. To investigate the subject in more detail, various types of modeling and calculations are increasingly used [269][270][271][272][273]. For example, the elastic properties of HA appeared to be significantly affected by the presence of vacancies, which softened HA via reducing its elastic modules [273]. In addition, a considerable anisotropy in the stress–strain behavior of the perfect HA crystals was found by ab initio calculations [270]. The crystals appeared to be brittle for tension along the z-axis with the maximum stress of ~9.6 GPa at 10% strain. Furthermore, the structural analysis of the HA crystal under various stages of tensile strain revealed that the deformation behavior manifested itself mainly in the rotation of PO4 tetrahedrons with concomitant movements of both the columnar and axial Ca ions [270]. Data for single crystals are also available [274]. Vickers hardness (a measure of the resistance to permanent indentation) of dense HA bioceramics is within 3–7 GPa, while the Poisson’s ratio (the ratio of the contraction or transverse strain to the extension or axial strain) for HA is about 0.27, which is close to that of bones (~0.3). At temperatures within 1000–1100 °C, dense HA bioceramics were found to exhibit superplasticity with a deformation mechanism based on grain boundary sliding [275][276][277]. Furthermore, both wear resistance and friction coefficient of dense HA bioceramics are comparable to those of dental enamel [254].
Due to a high brittleness (associated with a low crack resistance), the biomedical applications of CaPO4 bioceramics are focused on production of non-load-bearing implants, such as pieces for middle ear surgery, filling of bone defects in oral or orthopedic surgery, and coating of dental implants and metallic prosthesis (see below) [14][278][279]. Therefore, methods are continuously sought to improve the reliability of CaPO4 bioceramics. Namely, the mechanical properties of sintered bioceramics might be improved by changing the morphology of the initial CaPO4 [280]. In addition, diverse reinforcements (ceramics, metals, or polymers) have been applied to manufacture various biocomposites and hybrid biomaterials [281], but that is another story. However, successful hybrid formulations consisting of CaPO4 only [282][283][284][285][286][287][288][289] are within the scope. Namely, bulk HA bioceramics might be reinforced by HA whiskers [283][284][285][286][287]. Furthermore, various biphasic apatite/TCP formulations were tested [282][288][289] and, for example, a superior superplasticity of HA/β-TCP biocomposites to HA bioceramics was detected [288].
Another method to improve the mechanical properties of CaPO4 bioceramics is to cover the items by polymeric coatings [290][291][292] or infiltrate porous structures by polymers [293][294][295]; however, this is another topic. Other approaches are also possible [86]. Further details on the mechanical properties of CaPO4 bioceramics are available elsewhere [252][254][296], where interested readers are referred.

2.2. Electric/Dielectric and Piezoelectric Properties

Recently, an interest in both electric/dielectric [223][297][298][299][300][301][302][303][304][305][306][307][308][309] and piezoelectric [310][311] properties of CaPO4 bioceramics has been expressed. In addition, some types of CaPO4 bioceramics (namely, HA) appear to be electrets [312][313]. An electret is a dielectric material that has a quasi-permanent electric charge or dipole polarization. An electret generates internal and external electric fields, and is the electrostatic equivalent of a permanent magnet [314]. For example, a surface ionic conductivity of both porous and dense HA bioceramics was examined for humidity sensor applications, since the room temperature conductivity was influenced by relative humidity [298]. Namely, the ionic conductivity of HA is a subject of research for its possible use as a gas sensor for alcohol [299], carbon dioxide [297][306], or carbon monoxide [302]. Electric measurements were also used as a characterization tool to study the evolution of microstructure in HA bioceramics [300]. More to the point, the dielectric properties of HA were examined to understand its decomposition to β-TCP [299]. In the case of CDHA, the electric properties, in terms of ionic conductivity, were found to increase after compression of the samples at 15 t/cm2, which was attributed to establishment of some order within the apatitic network [301]. The conductivity mechanism of CDHA appeared to be multiple [304]. Furthermore, there are attempts to develop HA and/or CDHA electrets for biomedical utilization [303][312][313].
The electric properties of CaPO4 bioceramics appear to influence their biomedical applications. For example, there is an interest in polarization of HA bioceramics to generate a surface charge by the applying a constant DC electric field of 0.5–10.0 kV/cm at elevated temperatures (300–1000 °C) to samples previously sintered at ~ 1000–1250 °C for ~2 h. This technique is called thermally stimulated polarization and its results indicated that the polarization effects were a consequence of electrical dipoles associated with the formation of defects inside crystal grains, such as thermally-induced OH− vacancies, and of the space charge polarization that originated in the grain boundaries [315][316][317]. The presence of surface charges on HA was shown to have a significant effect on both in vitro and in vivo crystallization of biological apatite [318][319][320][321][322][323][324], as well as on an ability to adsorb various types of phosphate ions [317]. Furthermore, a growth of both biomimetic CaPO4 and bones was found to be accelerated on negatively charged surfaces and decelerated at positively charged surfaces [322][323][324][325][326][327][328][329][330][331]. A similar effect was found for adsorption of bovine serum albumin [332]. In addition, the electric polarization of CaPO4 was found to accelerate a cytoskeleton reorganization of osteoblast-like cells [333][334][335][336], extend bioactivity [337], enhance bone ingrowth through the pores of porous implants [338], and influence the cell activity [339][340]. The positive effect of electric polarization was found for carbonated apatite as well [341]. There is an interesting study on the interaction of a blood coagulation factor on electrically polarized HA surfaces [342]. Further details on the electric properties of CaPO4-based bioceramics are available in the literature [223][307][308][313].

2.3. Possible Transparency

Single crystals of all types of CaPO4 are optically transparent for the visible light. As bioceramics of CaPO4 have a polycrystalline nature with a random orientation of big amounts of small crystals, it is opaque and of white color, unless colored dopants have been added. However, in some cases, a transparency is convenient to provide some essential advantages (e.g., to enable direct viewing of living cells, their attachment, spreading, proliferation, and osteogenic differentiation cascade in a transmitted light). Thus, transparent CaPO4 bioceramics [343] have been prepared and investigated [28][112][114][275][343][344][345][346][347][348][349][350][351][352][353][354][355][356][357]. They can exhibit an optical transmittance of ~66% at a wavelength of 645 nm [355]. The preparation techniques include a hot isostatic pressing [28][112][114][356], an ambient-pressure sintering [350], a gel casting coupled with a low-temperature sintering [351][354], and a pulse electric current sintering [352], as well as spark plasma [275][344][345][346][347][348][349][358][359][360] and flash [229][230] sintering techniques. Fully dense, transparent CaPO4 bioceramics are obtained at temperatures above ~800 °C. Depending on the preparation technique, the transparent bioceramics have a uniform grain size ranging from ~67 nm [28] to ~250 μm [351] and are always pore-free. Furthermore, translucent CaPO4 bioceramics are also known [28][192][361][362][363]. Concerning possible biomedical applications, the optically transparent in visible light CaPO4 bioceramics can be useful for direct viewing of other objects, such as cells, in some specific experiments [353]. In addition, the transparency for laser light CaPO4 bioceramics may appear to be convenient for minimal invasive surgery by allowing passing the laser beam through it to treat the injured tissues located underneath. However, due to a lack of both porosity and the necessity to have see-through implants inside the body, the transparent and translucent forms of CaPO4 bioceramics will hardly be extensively used in medicine, except for the aforementioned cases and possible eye implants.

2.4. Porosity

Porosity is defined as a percentage of voids in solids, and this morphological property is independent of the material. The surface area of porous bodies is much higher, which guarantees a good mechanical fixation in addition to providing sites on the surface that allow chemical bonding between the bioceramics and bones [364]. Furthermore, a porous material may have both closed (isolated) pores and open (interconnected) pores. The latter look similar to tunnels and are accessible by gases, liquids, and particulate suspensions [365]. The open-cell nature of porous materials (also known as reticulated materials) is a unique characteristic essential in many applications. In addition, pore dimensions are also important. Namely, the dimensions of open pores are directly related to bone formation, since such pores grant both the surface and space for cell adhesion and bone ingrowth [366][367][368]. On the other hand, pore interconnection provides the ways for cell distribution and migration, and it allows an efficient in vivo blood vessel formation suitable for sustaining bone tissue neo-formation and possibly remodeling [54][338][369][370][371][372][373]. Namely, porous CaPO4 bioceramics are colonized easily by cells and bone tissues [369][372][374][375][376][377][378][379]. Therefore, interconnecting macroporosity (pore size > 100 μm) [25][364][369][380][381] is intentionally introduced in solid bioceramics. Calcining of natural bones and teeth appears to be the simplest way to prepare porous CaPO4 bioceramics [4][5][6][7][8][9][10][11]. In addition, macroporosity might be formed artificially due to a release of various easily removable compounds and, for that reason, incorporation of pore-creating additives (porogens) is the most popular technique to create macroporosity. The porogens are crystals, particles, or fibers of either volatile (they evolve gases at elevated temperatures) or soluble substances. The popular examples comprise paraffin [382][383][384], naphthalene [246][385][386][387], sucrose [388][389], NaHCO3 [390][391][392], NaCl [393][394], polymethylmethacrylate [395][396][397][398], hydrogen peroxide [399][400][401][402], cellulose [403], and its derivatives [16]. Several other compounds [261][404][405][406][407][408][409][410][411], including carbon nanotubes [412], might be used as porogens as well. The ideal porogen should be nontoxic and be removed at ambient temperature, thereby allowing the bioceramic/porogen mixture to be injected directly into a defect site and allowing the scaffold to fit the defect [413].
Many other techniques, such as replication of polymer foams by impregnation [147][148][149][152][414][415][416][417][418], various types of casting [130][131][135][137][402][419][420][421][422][423][424][425][426][427], suspension foaming [42], surfactant washing [428], microemulsions [429][430], and ice templating [431][432][433][434], as well as many other approaches [71][127][395][435][436][437][438][439][440][441][442][443][444][445][446][447][448][449][450][451][452][453][454][455][456][457][458][459][460][461][462], have been applied to fabricate porous CaPO4 bioceramics. Some of them are summarized [413]. In addition, both natural CaCO3 porous materials, such as coral skeletons [463][464], shells [464][465], and even wood [466], as well as artificially prepared ones [467], can be converted into porous CaPO4 under the hydrothermal conditions (250 °C, 24–48 h) with the microstructure undamaged. Porous HA bioceramics can also be obtained by hydrothermal hot pressing. This technique allows solidification of the HA powder at 100–300 °C (30 MPa, 2 h) [238]. In another approach, bi-continuous water-filled microemulsions are used as preorganized systems for the fabrication of needle-like frameworks of crystalline HA (2 °C, 3 weeks) [429][430]. In addition, porous CaPO4 might be prepared by a combination of gel casting and foam burn out methods [172][174], as well as by hardening of the self-setting formulations [383][384][391][392][393][394][454]. Lithography was used to print a polymeric material, followed by packing with HA and sintering [441]. Hot pressing was applied as well [468][469]. More to the point, an HA suspension can be cast into a porous CaCO3 skeleton, which is then dissolved, leaving a porous network [437]. A 3D periodic macroporous frame of HA was fabricated via a template-assisted colloidal processing technique [443][446]. In addition, porous HA bioceramics might be prepared by using different starting HA powders and sintering at various temperatures by a pressureless sintering [439]. Porous bioceramics with an improved strength might be fabricated from CaPO4 fibers or whiskers. In general, fibrous porous materials are known to exhibit an improved strength due to fiber interlocking, crack deflection, and/or pullout [470]. Namely, porous bioceramics with well-controlled open pores were processed by sintering of fibrous HA particles [438]. In another approach, porosity was achieved by firing apatite-fiber compacts mixed with carbon beads and agar. By varying the compaction pressure, firing temperature and carbon/HA ratio, the total porosity was controlled in the ranges from ~40% to ~85% [16]. Finally, a superporous (~85% porosity) HA bioceramic was developed as well [449][451][452]. Additional information on the processing routes to produce porous ceramics can be found in the literature [471].
Bioceramic microporosity (pore size < 10 μm), which is defined by its capacity to be impregnated by biological fluids [472], results from the sintering process, while the pore dimensions mainly depend on the material composition, thermal cycle, and sintering time. The microporosity provides both a greater surface area for protein adsorption and increased ionic solubility. For example, embedded osteocytes distributed throughout microporous rods might form a mechanosensory network, which would not be possible in scaffolds without microporosity [473][474]. CaPO4 bioceramics with nanodimensional (<100 nm) pores might be fabricated as well [475][476][477][478][479]. It is important to stress that differences in porogens usually influence the bioceramics’ macroporosity, while differences in sintering temperatures and conditions affect the percentage of microporosity. Usually, the higher the sintering temperature, the lower both the microporosity content and the specific surface area of bioceramics. Namely, HA bioceramics sintered at ~1200 °C show significantly less microporosity and a dramatic change in crystal sizes, if compared with those sintered at ~1050 °C [480]. Furthermore, the average shape of pores was found to transform from strongly oblate to round at higher sintering temperatures [481]. The total porosity (macroporosity + microporosity) of CaPO4 bioceramics was reported to be ~70% [482] or even ~85% [449][451][452] of the entire volume. In the case of coralline HA or bovine-derived apatites, the porosity of the original biologic material (coral or bovine bone) is usually preserved during processing [483]. To finalize the production topic, creation of the desired porosity in CaPO4 bioceramics is a rather complicated engineering task and interested readers are referred to the additional publications on the subject [261][368][453][484][485][486][487][488][489].
Regarding the biomedical importance of porosity, studies revealed that increasing of both the specific surface area and pore volume of bioceramics might greatly accelerate the in vivo process of apatite deposition and, therefore, enhance the bone-forming bioactivity. More importantly, a precise control over the porosity, pore dimensions, and internal pore architecture of bioceramics on different length scales is essential for understanding the structure–bioactivity relationship and the rational design of better bone-forming biomaterials [487][490][491]. Namely, in antibiotic charging experiments, CaPO4 bioceramics with nanodimensional (<100 nm) pores showed a much higher charging capacity (1621 μg/g) than those of commercially available CaPO4 (100 μg/g), which did not contain nanodimensional porosity [485]. In other experiments, porous blocks of HA were found to be viable carriers with sustained release profiles for drugs [492] and antibiotics over 12 days [493] and 12 weeks [494], respectively. Unfortunately, porosity significantly decreases the strength of implants [257][262][296]. Thus, porous CaPO4 implants cannot be loaded and are used to fill only small bone defects; however, their strength increases gradually when bones ingrow into the porous network of CaPO4 implants [52][495][496][497][498]. For example, bending strengths of 4–60 MPa for porous HA implants filled with 50%–60% of cortical bone were reported [495], while in another study an ingrown bone increased strength of porous HA bioceramics by a factor of three to four [497].
Unfortunately, the biomedical effects of bioceramics’ porosity are not straightforward. For example, the in vivo response of CaPO4 to different porosity was investigated, and a hardly any effect of macropore dimensions (~150, ~260, ~510, and ~1220 μm) was observed [499]. In another study, a greater differentiation of mesenchymal stem cells was observed when cultured on ~200 μm pore size HA scaffolds when compared to those on ~500 μm pore size HA [500]. The latter finding was attributed to the fact that a higher pore volume in ~500 μm macropore scaffolds might contribute to a lack of cell confluency, leading to the cells proliferating before beginning differentiation. In addition, the researchers hypothesized that bioceramics having less than the optimal pore dimensions induced quiescence in differentiated osteoblasts due to reduced cell confluency [500]. In still another study, the use of BCP (HA/TCP = 65/35 wt.%) scaffolds with cubic pores of ~500 μm resulted in the highest bone formation compared with the scaffolds with lower (~100 μm) or higher (~1000 μm) pore sizes [501]. Furthermore, CaPO4 bioceramics with greater strut porosity appeared to be more osteoinductive [502]. As early as 1979, Holmes suggested that the optimal pore range was 200–400 μm with the average human osteon size of ~223 μm [503]. In 1997, Tsurga and coworkers implied that the optimal pore size of bioceramics that supported ectopic bone formation was 300–400 μm [504]. Thus, there is no need to create CaPO4 bioceramics with very big pores; however, the pores must be interconnected [370][380][381][505]. Interconnectivity governs a depth of cells or tissue penetration into the porous bioceramics, and it allows development of blood vessels required for new bone nourishing and wastes removal [506][507]. Nevertheless, the total porosity of implanted bioceramics appears to be important. For example, 60% porous β-TCP granules achieved a higher bone fusion rate than 75% porous β-TCP granules in lumbar posterolateral fusion [473].
More details on the importance of CaPO4 bioceramics porosity on bone regeneration are available in a topical research [508].

3. Biomedical Applications

Since Levitt et al. described a method of preparing FA bioceramics and suggested their possible use in medical applications in 1969 [509], CaPO4 bioceramics have been widely tested for clinical applications. Namely, over 400 forms, compositions, and trademarks are currently either in use or under consideration in many areas of orthopedics and dentistry [510], with even more in development. In addition, various formulations containing demineralized bone matrix (commonly abbreviated as DBM) are produced for bone grafting. For example, bulk materials, available in dense and porous forms, are used for alveolar ridge augmentation, immediate tooth replacement, and maxillofacial reconstruction [511][512]. Other examples comprise burr-hole buttons [513][514], cosmetic (nonfunctional) eye replacements such as Bio-Eye® [515][516][517][518][519][520], increment of the hearing ossicles [521][522][523], and spine fusion [524][525][526][527], as well as repair of bone [51][528][529], craniofacial [530], and dental [531] defects. In order to permit growth of new bone into defects, a suitable bioresorbable material should fill these defects. Otherwise, ingrowth of fibrous tissue might prevent bone formation within the defects.
In spite of the aforementioned serious mechanical limitations (see Section 4.1. Mechanical Properties), bioceramics of CaPO4 are available in various physical forms: powders, particles, granules (or granulates), dense blocks, porous scaffolds, self-setting formulations, implant coatings, and composite components of different origin (natural, biological, or synthetic), often with specific shapes, such as implants, prostheses, or prosthetic devices. In addition, CaPO4 are also applied as nonhardening injectable formulations [532][533][534][535][536][537] and pastes [537][538][539][540][541]. Generally, they consist of a mixture of CaPO4 powders or granules and a “glue”, which can be a highly viscous hydrogel. More to the point, custom-designed shapes such as wedges for tibial opening osteotomy, cones for spine and knee, and inserts for vertebral cage fusion are also available [482].
One should note that among the existing CaPO4, only certain compounds are useful for biomedical applications, because those having a Ca/P ionic ratio less than 1 are not suitable for implantation due to their high solubility and acidity. Furthermore, due to its basicity, TTCP alone cannot be suitable either. Nevertheless, researchers try [72]. In addition, to simplify biomedical applications, these “of little use” CaPO4 can be successfully combined with either other types of CaPO4 or other chemicals.

3.1. Self-Setting (Self-Hardening) Formulations

The need for bioceramics for minimal invasive surgery has induced the concept of self-setting (or self-hardening) formulations consisting of CaPO4 only to be applied as injectable and/or moldable bone substitutes [43][44][55][441][542]. After hardening, they form bulk CaPO4 bioceramics. In addition, there are reinforced formulations that, in a certain sense, might be defined as CaPO4 concretes [43]. Furthermore, self-setting formulations able to produce porous bulk CaPO4 bioceramics are also available [383][384][391][392][393][394][441][454][542][543][544][545].
All types of the self-setting CaPO4 formulations belong to low-temperature bioceramics. They are divided into two major groups. The first one is a dry mixture of two different types of CaPO4 (a basic one and an acidic one), in which, after being wetted, the setting reaction occurs according to an acid–base reaction. The second group contains only one CaPO4, such as ACP with Ca/P molar ratio within 1.50–1.67 or α-TCP: both of them form CDHA upon contact with an aqueous solution [43][55]. Chemically, setting (= hardening, curing) is due to the succession of dissolution and precipitation reactions. Mechanically, it results from crystal entanglement and intergrowth [546]. By influencing dimensions of forming CaPO4 crystals, it is possible to influence the mechanical properties of the hardened bulk bioceramics [547]. Sometimes, the self-set formulations are sintered to prepare high-temperature CaPO4 bioceramics [548]. Despite a large number of initial compositions, all types of self-setting CaPO4 formulations can form three products only: CDHA, DCPD, and, rarely, DCPA [43][44][55][441][542]. Special research on the topic are available in [43][44][548], where interested readers are referred for further details.

3.2. CaPO4 Deposits (Coatings, Films, and Layers)

For many years, the clinical application of CaPO4-based bioceramics has been largely limited to non-load-bearing parts of the skeleton due to their inferior mechanical properties. Therefore, materials with better mechanical properties appear to be necessary. For example, metallic implants are encountered in endoprostheses (total hip joint replacements) and artificial teeth sockets. As metals do not undergo bone bonding, i.e., they do not form a mechanically stable link between the implant and bone tissue, methods have been sought to improve contacts at the interface. One major method is to coat metals with CaPO4, which enables bonding ability between the metal and the bone [107][117][321][549][550][551].
A number of factors influence the properties of CaPO4 deposits (coatings, films, and layers). They include thickness (this will influence coating adhesion and fixation—the agreed optimum now seems to be within 50–100 µm), crystallinity (this affects the dissolution and biological behavior), phase and chemical purity, porosity, and adhesion. The coated implants combine the surface biocompatibility and bioactivity of CaPO4 with the core strength of strong substrates. Moreover, CaPO4 deposits decrease a release of potentially hazardous chemicals from the core implant and shield the substrate surface from environmental attack. In the case of porous implants, the CaPO4-coated surface enhances bone ingrowth into the pores [254]. The clinical results for CaPO4-deposited implants reveal that they have much longer lifetimes after implantation than uncoated devices and they are found to be particularly beneficial for younger patients. Further details on this topic are available in the special research [549][550][551].

3.3. Functionally Graded Bioceramics

In general, functionally gradient materials (FGMs) are defined as materials having either compositional or structural gradient from their surface to the interior. The idea of FGMs allows one device to possess two different properties. One of the most important combinations for the biomedical field is that of mechanical strength and biocompatibility. Namely, only surface properties govern a biocompatibility of the entire device. In contrast, the strongest material determines the mechanical strength of the entire device. Although this subject belongs to the previous section on coatings, films, and layers, in a certain sense, all types of implants covered by CaPO4 might be also considered as FGMs.
Within the scope, functionally graded bioceramics consisting of CaPO4 are considered and discussed only. Such formulations have been developed [395][423][426][486][552][553][554][555][556][557][558][559][560][561][562]. For example, dense sintered bodies with gradual compositional changes from α-TCP to HA were prepared by sintering diamond-coated HA compacts at 1280 °C under a reduced pressure, followed by heating under atmospheric conditions [552]. The content of α-TCP gradually decreased, while the content of HA increased with increasing depth from the surface. This functionally gradient bioceramic consisting of HA core and α-TCP surface showed potential value as a bone-substituting biomaterial [552]. Two types of functionally gradient FA/β-TCP biocomposites were prepared in another study [553].
In addition, it is well known that a bone cross-section from cancellous to cortical bone is nonuniform in porosity and pore dimensions. Thus, in various attempts to mimic the porous structure of bones, CaPO4 bioceramics with graded porosity have been fabricated [365][395][412][423][426][486][552][553][554][555]. For example, graded porous CaPO4 bioceramics can be produced by means of tape casting and lamination. Other manufacturing techniques, such as a compression molding process followed by impregnation and firing, are known as well [365]. In the first method, an HA slurry was mixed with a pore former. The mixed slurry was then cast into a tape. Using the same method, different tapes with different pore former sizes were prepared individually. The different tape layers were then laminated together. Firing was then performed to remove the pore formers and sinter the HA particle compacts, resulting in graded porous bioceramics [555]. This method was also used to prepare graded porous HA with a dense part (core or layer) in order to improve the mechanical strength, as dense ceramics are much stronger than porous ceramics. However, as in the pressure infiltration of mixed particles, this multiple tape casting also has the problem of poor connectivity of pores, although the pore size and the porosity are relatively easy to control. Furthermore, the lamination step also introduces additional discontinuity of the porosity on the interfaces between the stacked layers.
Since diverse biomedical applications require different configurations and shapes, the graded (or gradient) porous bioceramics can be grouped according to both the overall shape and the structural configuration [365]. The basic shapes include rectangular blocks and cylinders (or disks). For the cylindrical shape, there are configurations of dense core–porous layer, less porous core–more porous layer, dense layer–porous core, and less porous layer–more porous core. For the rectangular shape, in the gradient direction, i.e., the direction with varying porosity, pore size, or composition, there are configurations of porous top–dense bottom (same as porous bottom–dense top), porous top–dense center–porous bottom, dense top–porous center–dense bottom, etc. Concerning biomedical applications, a dense core–porous layer structure is suitable for implants of a high mechanical strength and with bone ingrowth for stabilization, whereas a less porous layer–more porous core configuration can be used for drug delivery systems. Furthermore, a porous top –dense bottom structure can be shaped into implants of articulate surfaces for wear resistance and with porous ends for bone ingrowth fixation, while a dense top–porous center–dense bottom arrangement mimics the structure of head skull. Further details on bioceramics with graded porosity can be found in the literature [365].

4. Non-Biomedical Applications of CaPO4

Due to their strong adsorption ability, surface acidity or basicity, and ion exchange abilities, some types of CaPO4 possess a catalytic activity [563][564][565][566][567][568][569][570][571][572][573][574][575]. As seen from the references, CaPO4 are able to catalyze oxidation and reduction reactions, as well as formation of C–C bonds. Namely, the application in oxidation reactions mainly includes oxidation of alcohol and dehydrogenation of hydrocarbons, while the reduction reactions include hydrogenolysis and hydrogenation. The formation of C–C bonds mainly comprises Claisen–Schmidt and Knoevenagel condensation reactions, Michael addition reaction, as well as Friedel–Crafts, Heck, Diels–Alder, and aldol reactions [570].
In addition, due to the chemical similarity to the inorganic part of mammalian calcified tissues, CaPO4 powders appear to be good solid carriers for chromatography of biological substances. Namely, high-value biological materials such as recombinant proteins, therapeutic antibodies, and nucleic acids are separated and purified [576][577][578][579][580][581][582]. Furthermore, some types of CaPO4 are used as a component of various sensors [297][298][302][303][306][583][584][585][586][587]. Finally, CaPO4 ceramics appear to be good adsorbents of fluorides [588]; however, since these subjects are almost irrelevant to bioceramics, they are not detailed further. Additional details and examples are available elsewhere [589].

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