In order to fabricate scaffolds 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 137138 and rapid prototyping 139140141 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 the 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 3D CAD computer model 142143144145146. 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 147148, selective laser sintering 149150151152153154155, laser cladding 156157158159, 3D printing and/or plotting 86160161162163164165166167168169170171172173174, solid freeform fabrication 175176177178179180181182183 and stereolithography 184185186187. 3D printing and/or plotting of the CaPO₄ -based self-setting formulations could be performed as well 172173. In the specific case of ceramic scaffolds, a sintering step is usually applied after printing the green bodies. Furthermore, a thermal printing process of melted CaPO₄ has been proposed 188, while, in some cases, laser processing might be applied as well 189190. A schematic of 3D printing technique, as well as some 3D printed items are shown in Figure 410. A custom-made implant of actual dimensions would reduce the time it takes to perform the medical implantation procedure and subsequently lower the risk to the patient. Another advantage of a pre-fabricated, exact-fitting implant is that it can be used more effectively and applied directly to the damaged site rather than a replacement, which is formulated during surgery from a paste or granular material 176190191192.

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 scaffolds, e.g., whether they are for a hard-tissue replacement or an integration of the device within the surrounding tissues. In general, three types of the 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 193194195196. 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 b ioceramics and scaffolds. The methods include uniaxial compaction 197198199, isostatic pressing (cold or hot) 100200201202203204205206207, granulation 208209210211212213214, loose packing 215, slip casting 88216217218219220221, gel casting 184222223224225226227228229230, pressure mold forming 231232, injection molding 233234235, polymer replication 236237238239240241242243, ultrasonic machining 244, extrusion 245246247248249250251, slurry dipping and spraying 252. Depending on the fabrication technique used, various parameters such as solid loading of the ceramic slurry, type and amount of additives (binders, surfactants, dispersants, etc.), temperature, etc. should be optimized to maximize the mechanical strength of the scaffolds. In addition, to form ceramic sheets from slurries, tape casting 225253254255256257, doctor blade 258 and colander methods can be employed 193194195196. Furthermore, 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 259. Various combinations of several techniques are also possible 90225260261262. More to the point, some of these processes might be performed under the electromagnetic field, which helps crystal aligning 217220263264265266. Finally, the prepared CaPO₄ bioceramics might be subjected by additional treatments (e.g., chemical, thermal and/or hydrothermal ones) to convert one type of CaPO₄ into another one 243.

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 267. 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 193194195196. 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 196. 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 scaffolds 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 196. 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 scaffolds, machine finishing might be essential 144193268269. Unfortunately, cutting tools developed for metals are usually useless for bioceramics due to their fragility; therefore, grinding and polishing appear to be the convenient finishing techniques 144193. In addition, the surface of scaffolds might be modified by various supplementary treatments 270.

A sintering (or firing) procedure appears to be of a great importance to manufacture bioceramic scaffolds with the required mechanical properties. Usually, this stage 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 193194195196. 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 271. In the majority cases, sintering allows a structure to retain its shape. However, this process might be accompanied by a considerable degree of shrinkage 272273274, 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 275.

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 276. Furthermore, strong chemical bonds are formed among the particles and loosely compacted green bodies are hardened to denser materials 193194195196. Further knowledge on the ceramic sintering process might be found elsewhere 277.

In the case of CaPO₄, the earliest paper on their sintering was published in 1971 278. Since then, numerous papers on this subject were published and several specific processes were 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., 2HPO₄^2- → P₂O₇^4- + H₂O↑) takes place. Besides, sintering causes toughening 79, densification 80279, partial dehydroxylation (in the case of HA) 80, grain growth 276280, as well as it increases the mechanical strength 281282283. The latter events are due to presence of air and other gases filling gaps among the particles of un-sintered 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 a biologically formed apatites was investigated 284285 and the obtained products were characterized 286287. In all cases, the numerical value of 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 cases, CaPO₄ with Ca/P ratio < 1.5 (Table 1) are not sintered, since these compounds are thermally unstable, while sintering of non-stoichiometric CaPO₄ (CDHA and ACP) always leads to their transformation into various types of biphasic, triphasic and multiphase formulations 92.

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 288. 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 288. To enhance sinterability of CaPO₄, a variety of sintering additives might be added 289290291292.

Solid-state pressureless sintering is the simplest procedure. For example, HA scaffolds 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 125126127128129, 293,294,295,296. The decomposition temperature of HA 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 from the decomposition. On the other hand, a presence of water in the sintering atmosphere was reported to inhibit densification of HA and accelerated grain growth 297. Unexpectedly, an application of a magnetic field during sintering was found to influence the growth of HA grains 280. 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 298299300.

Since grain growth occurs mainly during the final stage of sintering, to avoid this, a new method called ‘‘two-step sintering’’ (TSS) was proposed 301. 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 9199302303304305306. 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 the fracture toughness of samples from 0.98 ± 0.12 to 1.92 ± 0.20 MPa m^1/2. In addition, due to the lower second step sintering temperature, no HA phase decomposition was detected in TSS method 302.

Hot pressing 300307308309310311312313, hot isostatic pressing 100200205207 or hot pressing with post-sintering 314315 processes make it possible to decrease a temperature of the densification process, diminish the grain size, as well as achieve higher densities. This leads to finer microstructures, higher thermal stability and subsequently better mechanical properties of CaPO₄ scaffolds. Both microwave 316317318319320321322323324325 and spark plasma 82116326327328329330331332333334335 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 has been developed to fabricate OCP 117, CDHA 336, HA/β-TCP 310 and HA 311312313314 bioceramics with neither thermal dehydration nor thermal decomposition. Further details on the sintering and firing processes of CaPO₄ bioceramics are available in literature 127338339.

To conclude this section, one should mention that the sintering stage is not always necessary. For example, CaPO₄ -based scaffolds with the reasonable mechanical properties might be prepared by means of self-setting (self-hardening) formulations 61. Furthermore, the reader’s attention is paid on an excellent review on various ceramic manufacturing techniques 340.

The modern generation of biomedical materials should stimulate the body’s own self-repairing abilities 341. 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 up to three times body weight (3,000 N) and peak loads experienced during jumping can be as high as 10 times 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 342. 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 194. 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 one with larger grain sizes. Thus, in general, the strength for ceramics is proportional to the inverse square root of the grain sizes 343. 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 a slow crack growth 344. Accordingly, from the mechanical point of view, CaPO₄ scaffolds appear to be brittle polycrystalline materials for which the mechanical properties are governed by crystallinity, grain size, grain boundaries, porosity and composition 345. Thus, it possesses poor mechanical properties (for instance, a low impact and fracture resistances) that do not allow CaPO₄ scaffolds to be used in load-bearing areas, such as artificial teeth or bones. For example, fracture toughness (this is a property, which 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·m^1/2 346 (human bone: 2 – 12 MPa·m^1/2). It decreases exponentially with the porosity increasing 347. Generally, fracture toughness increases with grain size decreasing. However, in some materials, especially non-cubic 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 authors 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·m^1/2 at ~ 0.4 µm 307. Fracture energy of HA bioceramics is in the range of 2.3 – 20 J/m², while the Weibull modulus (it is 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 to a low reliability of HA implants 348. Porosity has a great influence on the Weibull modulus 349350. In addition, that the reliability of HA scaffolds was found to depend on deformation mode (bending or compression), along with pore size and pore size distribution: a reliability was higher for smaller average pore sizes in bending but lower for smaller pore sizes in compression 351. Interestingly that 3 peaks of internal friction were found at temperatures 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 352. The differences in internal friction values were also found between HA and TCP 353.

Bending, compressive and tensile strengths of dense HA bioceramics are in the ranges of 38 – 250 MPa, 120 – 900 MPa and 38 – 300 MPa, respectively. Similar values for porous HA scaffolds are substantially lower: 2 – 11 MPa, 2 – 100 MPa and ~ 3 MPa, respectively 348. 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, as well as they are 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 decreases suddenly when Ca/P > 1.67 348. Furthermore, strength decreases almost exponentially with porosity increasing 354355. However, by changing the pore geometry, it is possible to influence the strength of porous bioceramics. It is also worth mentioning that porous CaPO₄ scaffolds are considerably less fatigue resistant than dense bioceramics (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 a little effect on the fracture toughness of CaPO₄ bioceramics 345356. 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 357.

Young’s (or elastic) modulus of dense HA bioceramics is in the range of 35 – 120 GPa 358359, 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 360. Nevertheless, dense bulk compacts of HA have mechanical resistances of the order of 100 MPa versus ~ 300 MPa of human bones, diminishing drastically their resistances in the case of porous bulk compacts 361. Young’s modulus measured in bending is between 44 and 88 GPa. To investigate the subject in more details, various types of modeling and calculations are increasingly used 362363364365366. 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 366. In addition, a considerable anisotropy in the stress-strain behavior of the perfect HA crystals was found by ab initio calculations 363. 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 363. Data for single crystals are also available 367. Vickers hardness (that is a measure of the resistance to permanent indentation) of dense HA bioceramics is within 3 – 7 GPa, while the Poisson’s ratio (that is 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 was found to exhibit superplasticity with a deformation mechanism based on grain boundary sliding 332368369. Furthermore, both wear resistance and friction coefficient of dense HA bioceramics are comparable to those of dental enamel 348.

Due to a high brittleness (associated to 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, as well as coating of dental implants and metallic prosthesis (see below) 370371. Therefore, ways 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 CaPO₄372. In addition, diverse reinforcements (ceramics, metals or polymers) have been applied to manufacture various biocomposites and hybrid biomaterials 373, but that is another story. However, successful hybrid formulations consisted of CaPO₄ only 374375376377378379380381 are within the scope of this review. Namely, bulk HA bioceramics might be reinforced by HA whiskers 375376377378379. Furthermore, various biphasic apatite/TCP formulations were tested 374380381 and, for example, a superior superplasticity of HA/β-TCP biocomposites to HA bioceramics was detected 380.

Another approach to improve the mechanical properties of CaPO₄ bioceramics is to cover the items by polymeric coatings 382383384 or infiltrate porous structures by polymers 385386387; however, this is still other story. Further details on the mechanical properties of CaPO₄ bioceramics are available elsewhere 347348388, where the interested readers are referred to.

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 bioceramic scaffolds and bones 389. Furthermore, a porous material may have both closed (isolated) pores and open (interconnected) pores. The latter look like tunnels, which are accessible by gases, liquids and particulate suspensions 390. 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 391392393. On the other hand, pore interconnection provides the ways for cell distribution and migration, as well as it allows an efficient in vivo blood vessel formation suitable for sustaining bone tissue neo-formation and possibly remodeling 135394395396397398399. Explicitly, porous CaPO₄ scaffolds are colonized easily by cells and bone tissues 394398400401402403404405406407. Therefore, interconnecting macroporosity (pore size > 100 μm) 97389394408409 is intentionally introduced into solid scaffolds (Figure 5). Calcining of natural bones appears to be the simplest way to prepare porous CaPO₄ scaffolds 68697071727374. 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 410411412, naphthalene 345413414415, sucrose 416417, NaHCO₃418419420, NaCl 421422, polymethylmethacrylate 87423424425, hydrogen peroxide 426427428429430431, cellulose derivatives 77. Several other compounds 338355432433434435436437438439440441442443 might be used as porogens either. 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 444. Sintering particles, preferably spheres of equal size, is a similar way to generate porous 3D scaffolds of CaPO₄. However, pores resulting from this method are often irregular in size and shape and not fully interconnected with one another. Schematic drawings of various types of the ceramic porosity are shown in Figure 6445.

Many other techniques, such as replication of polymer foams by impregnation 236237238446,447,448,449,450 (Figure 5), various types of casting 218219225227431451452453454455456457458459, suspension foaming 114, surfactant washing 460, microemulsions 461462, ice templating 463464465466, as well as many other approaches 1281848788154467468469470471472473474475476477478479480481482483484485486487488489490491492493494495496497498499500501 have been applied to fabricate porous CaPO₄ scaffolds. Some of them have been summarized in Table 3444. In addition, both natural CaCO₃ porous materials, such as coral skeletons 502503 or shells 503504, and artificially prepared ones 505 can be converted into porous CaPO₄ under the hydrothermal conditions (250 °C, 24 – 48 h) with the microstructure undamaged. Porous HA scaffolds can also be obtained by hydrothermal hot pressing. This technique allows solidification of the HA powder at 100 – 300 °C (30 MPa, 2 h) 337. In another approach, bi-continuous water-filled microemulsions have been used as pre-organized systems for the fabrication of needle-like frameworks of crystalline HA (2 °C, 3 weeks) 461462. Besides, porous CaPO₄ might be prepared by a combination of gel casting and foam burn out methods 260262, as well as by hardening of the self-setting formulations 411412419420422432433490. Lithography was used to print a polymeric material, followed by packing with HA and sintering 471. Hot pressing was applied as well 308309. More to the point, a HA suspension can be cast into a porous CaCO₃ skeleton, which is then dissolved, leaving a porous network assisted colloidal processing technique 472473. In addition, porous HA scaffolds might be prepared by using different starting HA powders and sintering at various temperatures by a pressureless sintering 469. Porous scaffolds 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 27. Namely, porous scaffolds with well-controlled open pores was processed by sintering of fibrous HA particles 468. 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 % 77. Finally, a superporous (~ 85 % porosity) HA scaffolds was developed as well 60487488. Additional information on the processing routes to produce porous ceramics might be found in literature 506507.

Scaffold microporosity (pore size < 10 μm), which is defined by its capacity to be impregnated by biological fluids 508, 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 509510. CaPO₄ scaffolds with nanodimensional (< 100 nm) pores might be fabricated as well 195511512513514515. It is important to stress, that differences in porogens usually influence the scaffolds’ 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 the scaffolds. Namely, HA bioceramics sintered at ~ 1200 °C shows significantly less microporosity and a dramatic change in crystal sizes, if compared with that sintered at ~ 1050 °C (Figure 7) 516. Furthermore, the average shape of pores was found to transform from strongly oblate to round at higher sintering temperatures 517. The total porosity (macroporosity + microporosity) of CaPO₄ scaffolds was reported to be ~ 70 % 518 or even ~ 85 % 60487488 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 519. To finalize the production topic, creation of the desired porosity in CaPO₄ scaffolds is a rather complicated engineering task and the interested readers are referred to the additional publications on the subject 355393489520521522523524525526527528.

Regarding the biomedical importance of porosity, studies revealed that increasing of both the specific surface area and pore volume of scaffolds 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 the scaffolds on different length scales is essential for understanding of the structure-bioactivity relationship and the rational design of better bone-forming biomaterials 526529530. Namely, in antibiotic charging experiments, CaPO₄ scaffolds with nanodimensional (< 100 nm) pores showed a much higher charging capacity (1621 μg/g) than that of commercially available CaPO₄ (100 μg/g), which did not contain nanodimensional porosity 522. In other experiments, porous blocks of HA were found to be viable carriers with sustained release profiles for drugs 531 and antibiotics over 12 days 532 and 12 weeks 533, respectively. Unfortunately, porosity significantly decreases the strength of implants 348356388. 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 131534535536537. For example, bending strengths of 40 – 60 MPa for porous HA implants filled with 50 – 60 % of cortical bone were reported 534, while in another study an ingrown bone increased strength of porous HA scaffolds by a factor of 3 to 4 536.

Unfortunately, the biomedical effects of scaffolds’ porosity are not straightforward. For example, the in vivo response of CaPO₄ of different porosity was investigated and a hardly any effect of macropore dimensions (~ 150, ~ 260, ~ 510 and ~ 1220 μm) was observed 538. 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 539. 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. Besides, the authors hypothesized that scaffolds with a less than the optimal pore dimensions induced quiescence in differentiated osteoblasts due to reduced cell confluences 539. 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 scaffold with lower (~ 100 μm) or higher (~ 1000 μm) pore sizes 540. Furthermore, CaPO₄ scaffolds with greater strut porosity appeared to be more osteoinductive 541. Already in 1979, Holmes suggested that the optimal pore range was 200 – 400 μm with the average human osteon size of ~ 223 μm 542. In 1997, Tsurga and coworkers implied that the optimal pore size of scaffolds that supported ectopic bone formation was 300 – 400 μm 543. Thus, there is no need to create CaPO₄ scaffolds with very big pores; however, the pores must be interconnected 42408409544. Interconnectivity governs a depth of cells or tissue penetration into the porous scaffolds, as well as it allows development of blood vessels required for new bone nourishing and wastes removal 508545. Nevertheless, the total porosity of implanted scaffolds 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 509.

All interactions between implants and the surrounding tissues are dynamic processes. Water, dissolved ions, various biomolecules and cells surround the implant surface within initial few seconds after the implantation. It has been accepted that no foreign material placed inside a living body is completely compatible. The only substances that conform completely are those manufactured by the body itself (autogenous), while any other substance, which is recognized as foreign, initiates some types of reactions (a host-tissue response). The reactions occurring at the biomaterial/tissue interfaces lead to time-dependent changes in the surface characteristics of both the implanted biomaterials and the surrounding tissues 567.

In order to develop new scaffolds, it is necessary to understand the in vivo host responses. Like any other species, biomaterials and bioceramics react chemically with their environment and, ideally, they should neither induce any changes nor provoke undesired reactions in the neighboring or distant tissues. In general, living organisms can treat artificial implants as biotoxic (or bioincompatible 568), bioinert (or biostable 59), biotolerant (or biocompatible 568), bioactive and bioresorbable materials 566567568569. Biotoxic (e.g., alloys containing cadmium, vanadium, lead and other toxic elements) materials release to the body substances in toxic concentrations and/or trigger the formation of antigens that may cause immune reactions ranging from simple allergies to inflammation to septic rejection with the associated severe health consequences. They cause atrophy, pathological change or rejection of living tissue near the material as a result of chemical, galvanic or other processes. Bioinert (this term should be used with care, since it is clear that any material introduced into the physiological environment will induce a response. However, for the purposes of biomedical implants, the term can be defined as a minimal level of response from the host tissue), such as zirconia, alumina, carbon and titanium, as well as biotolerant (e.g., polymethylmethacrylate, titanium and Co-Cr alloy) materials do not release any toxic constituents but also do not show positive interaction with living tissue. They evoke a physiological response to form a fibrous capsule, thus, isolating the material from the body. In such cases, thickness of the layer of fibrous tissue separating the material from other tissues of an organism can serve as a measure of bioinertness. Generally, both bioactivity and bioresorbability phenomena are fine examples of chemical reactivity and CaPO₄ (both non-substituted and ion-substituted ones) fall into these two categories of bioceramics 566567568569. A bioactive material will dissolve slightly but promote formation of a surface layer of biological apatite before interfacing directly with the tissue at the atomic level, that result in formation of a direct chemical bonds to bones. Such implants provide a good stabilization for materials that are subject to mechanical loading. A bioresorbable material will dissolve over time (regardless of the mechanism leading to the material removal) and allow a newly formed tissue to grow into any surface irregularities but may not necessarily interface directly with the material. Consequently, the functions of bioresorbable materials are to participate in dynamic processes of formation and re-absorption occurring in bone tissues; thus, bioresorbable materials are used as scaffolds or filling spacers allowing to the tissues their infiltration and substitution 193338570571572.

It is important to stress, that a distinction between the bioactive and bioresorbable bioceramics might be associated with structural factors only. Namely, bioceramics made from non-porous, dense and highly crystalline HA behaves as a bioinert (but a bioactive) material and is retained in an organism for at least 5 – 7 years without noticeable changes (Figure 3 bottom), while a highly porous bioceramics of the same composition can be resorbed approximately within a year. Furthermore, submicron-sized HA powders are biodegraded even faster than the highly porous HA scaffolds. Other examples of bioresorbable materials comprise porous bioceramic scaffolds made of biphasic, triphasic or multiphasic CaPO₄ formulations 92 or bone grafts (dense or porous) made of CDHA 133, TCP 87573574 and/or ACP 434575. One must stress that at the beginning of 2000-s the concepts of bioactive and bioresorbable materials have been converged and bioactive materials are made bioresorbable, while bioresorbable ones are made bioactive 576.

Although in certain in vivo experiments inflammatory reactions were observed after implantation or injection of CaPO₄577578579580581582583584585586, the general conclusion on using CaPO₄ with Ca/P ionic ratio within 1.0 – 1.7 is that all types of implants (scaffolds of various porosities and structures, as well as, powders or granules) are not only nontoxic but also induce neither inflammatory nor foreign-body reactions 120587588. The biological response to implanted CaPO₄ scaffolds follows a similar cascade observed in fracture healing. This cascade includes a hematoma formation, inflammation, neovascularization, osteoclastic resorption and a new bone formation. An intermediate layer of fibrous tissue between the implants and bones has been never detected. Furthermore, CaPO₄ implants display the ability to directly bond to bones 566567568569. For further details, the interested readers are referred to a good review on cellular perspectives of bioceramic scaffolds for bone tissue engineering 444.

One should note that the aforementioned rare cases of the inflammatory reactions to CaPO₄ scaffolds were often caused by “other” reasons. For example, a high rate of wound inflammation occurred when highly porous HA scaffolds were used. In that particular case, the inflammation was explained by sharp implant edges, which irritated surrounding soft tissues 578. To avoid this, only rounded material should be used for implantation (Figure 10) 589. Another reason for inflammation produced by HA scaffolds could be due to micro movements of the implants, leading to simultaneous disruption of a large number of micro-vessels, which grow into the pores of scaffolds. This would immediately produce an inflammatory reaction. Additionally, problems could arise in clinical tests connected with migration of granules used for alveolar ridge augmentation, because it might be difficult to achieve a mechanical stability of implants at the implantation sites 578. Besides, presence of calcium pyrophosphate impurity might be the reason of inflammation 581. Additional details on inflammatory cell responses to CaPO₄ might be found in a special review on this topic 582.

Before recently, it was generally considered, that alone, any type of synthetic scaffolds possessed neither osteogenic (osteogenesis is the process of laying down new bone material by osteoblasts 590) nor osteoinductive (is the property of the material to induce bone formation de novo or ectopically (i.e., in non-bone forming sites) 590) properties and demonstrated a minimal immediate structural support. However, a number of reports have already shown the osteoinductive properties of certain types of CaPO₄ bioceramics 164516541591592593594595596597598599600601602603604605606607608609610 and the amount of such publications rapidly increases. For example, bone formation was found to occur in dog muscle inside porous CaPO₄ scaffolds with surface microporosity, while bone was not observed on the surface of dense bioceramics 595. Furthermore, implantation of porous β-TCP scaffolds appeared to induce bone formation in soft tissues of dogs, while no bone formation was detected in any α-TCP ones 592. More to the point, titanium implants coated by a microporous layer of OCP were found to induce ectopic bone formation in goat muscles, while a smooth layer of carbonated apatite on the same implants was not able to induce bone formation there 593594. In another study, β-TCP powder, biphasic (HA + β-TCP) powder and intact biphasic (HA + β-TCP) rods were implanted into leg muscles of mice and dorsal muscles of rabbits 601. One month and three months after implantation, samples were harvested for biological and histological analysis. New bone tissues were observed in 10 of 10 samples for β-TCP powder, 3 of 10 samples biphasic powder and 9 of 10 samples for intact biphasic rods at 3rd month in mice, but not in rabbits. The authors concluded that the chemical composition was the prerequisite in osteoinduction, while porosity contributed to more bone formation 601. Therefore, researchers have already discovered the ways to prepare osteoinductive CaPO₄ scaffolds.

Unfortunately, the underlying mechanism(s) leading to bone induction by synthetic materials remains largely unknown. Nevertheless, besides the specific genetic factors 599 and chosen animals 601, the dissolution/precipitation behavior of CaPO₄611, their particle size 609, microporosity 597601612613, physicochemical properties 595597, composition 601, the specific surface area 613, nanostructure 600, as well as the surface topography and geometry 596614615616617618 have been pointed out as the relevant parameters. A positive effect of increased microporosity on the ectopic bone formation could be both direct and indirect. Firstly, an increased microporosity is directly related to the changes in surface topography, i.e. increases a surface roughness, which affects the cellular differentiation 618. Secondly, an increased microporosity indirectly means a larger surface that is exposed to the body fluids leading to elevated dissolution/precipitation phenomena as compared to non-microporous surfaces. In addition, other hypotheses are also available. Namely, Reddi explained the apparent osteoinductive properties as an ability of particular bioceramics to concentrate bone growth factors, which are circulating in biological fluids, and those growth factors induce bone formation 614. Other researchers proposed a similar hypothesis that the intrinsic osteoinduction by CaPO₄ scaffolds was a result of adsorption of osteoinductive substances on their surface 596. Moreover, Ripamonti 615 and Kuboki et al., 616 independently postulated that the geometry of CaPO₄ bioceramics is a critical parameter in bone induction. Specifically, bone induction by CaPO₄ was never observed on flat surfaces. All osteoinductive cases were observed on either porous scaffolds or structures contained well-defined concavities. What’s more, bone formation was never observed on the peripheries of porous scaffolds and was always found inside the pores or concavities, aligning the surface 193. Some researchers speculated that a low oxygen tension in the central region of implants might provoke a dedifferentiation of pericytes from blood micro-vessels into osteoblasts 619. Finally but yet importantly, both nano-structured rough surfaces and a surface charge on implants were found to cause an asymmetrical division of the stem cells into osteoblasts, which is important for osteoinduction 612.

Nevertheless, to finalize this topic, it is worth citing a conclusion made by Boyan and Schwartz 620: “Synthetic materials are presently used routinely as osteoconductive bone graft substitutes, but before purely synthetic materials can be used to treat bone defects in humans where an osteoinductive agent is required, a more complete appreciation of the biology of bone regeneration is needed. An understanding is needed of how synthetic materials modulate the migration, attachment, proliferation and differentiation of mesenchymal stem cells, how cells on the surface of a material affect other progenitor cells in the peri-implant tissue, how vascular progenitors can be recruited and a neovasculature maintained, and how remodeling of newly formed bone can be controlled.” (p. 9).

Shortly after implantation, a healing process is initiated by compositional changes of the surrounding bio-fluids and adsorption of biomolecules. Following this, various types of cells reach the surface of CaPO₄ scaffolds and the adsorbed layer dictates the ways the cells respond. Further, a biodegradation (which can be envisioned as an in vivo process by which an implanted material breaks down into either simpler components or components of the smaller dimensions) of the implanted CaPO₄ scaffolds begin. This process can occur by three possible ways: 1) physical: due to abrasion, fracture and/or disintegration, 2) chemical: due to physicochemical dissolution of the implanted phases of CaPO₄ with a possibility of phase transformations into other phases of CaPO₄, as well as their precipitation and 3) biological: due to cellular activity (so called, bioresorption). In biological systems, all these processes take place simultaneously and/or in competition with each other. Since the existing CaPO₄ are differentiated by Ca/P ratio, basicity/acidity and solubility (Table 1), in the first instance, their degradation kinetics and mechanisms depend on the chosen type of CaPO₄621622. Since dissolution is a physical chemistry process, it is controlled by some factors, such as CaPO₄ solubility, surface area to volume ratio, local acidity, fluid convection and temperature. For HA and FA, the dissolution mechanism in acids has been described by a sequence of four successive chemical equations, in which several other CaPO₄, such as TCP, DCPD/DCPA and MCPM/MCPA, appear as virtual intermediate phases 623624.

With a few exceptions, dissolution rates of CaPO₄ are inversely proportional to the Ca/P ratio (except of TTCP), phase purity and crystalline size, as well as it is directly related to both the porosity and the surface area. In addition, phase transformations might occur with DCPA, DCPD, OCP, α-TCP, β-TCP and ACP because they are unstable in aqueous environment under the physiological conditions 625. Bioresorption is a biological process mediated by cells (mainly, osteoclasts and, in a lesser extent, macrophages) 626627. It depends on the response of cells to their environment. Osteoclasts attach firmly to the implant and dissolve CaPO₄ by secreting an enzyme carbonic anhydrase or any other acid, leading to a local pH drop to ~ 4 – 5 628. Formation of multiple spine-like crystals at the exposed areas of β-TCP was discovered 629. Furthermore, nanodimensional particles of CaPO₄ can also be phagocytosed by cells, i.e. they are incorporated into cytoplasm and thereafter dissolved by acid attack and/or enzymatic processes 630. A study is available 631, in which a comparison was made between the solubility and osteoclastic resorbability of 3 types of CaPO₄ (DCPA, ACP and HA) + β-calcium pyrophosphate (β-CPP) powders having the monodisperse particle size distributions. The authors discovered that with the exception of β-CPP, the difference in solubility among different calcium phosphates became neither mitigated nor reversed but augmented in the resorptive osteoclastic milieu. Namely, DCPA (the phase with the highest solubility) was resorbed more intensely than any other calcium phosphate, whereas HA (the phase with the lowest solubility) was resorbed the least. β-CPP became retained inside the cells for the longest period of time, indicating hindered digestion of only this particular type of calcium phosphate. Genesis of osteoclasts was found to be mildly hindered in the presence of HA, ACP and DCPA, but not in the presence of β-CPP. HA appeared to be the most viable compound with respect to the mitochondrial succinic dehydrogenase activity. The authors concluded that chemistry did have a direct effect on biology, while biology neither overrode nor reversed the chemical propensities of calcium phosphates with which it interacted, but rather augmented and took a direct advantage of them 631. Similar conclusions on both the resorbability and dissolution behavior of OCP, β-TCP and HA were made in another study 625. In addition, in vivo biodegradation of MCPA was found to be faster than that of bovine HA 632. Thus, one can conclude that in vivo biodegradation kinetics of CaPO₄ seems to correlate well with their solubility. Nevertheless, one must keep in mind that this is a very complicated combination of various non-equilibrium processes, occurring simultaneously and/or in competition with each other 633.

Strictly speaking, the processes happen in vitro do not necessarily represent the ones occurring in vivo and vice versa; nevertheless, in vitro experiments are widely performed. Usually, an in vitro biodegradation of CaPO₄ scaffolds is simulated by suspending them in a slightly acidic (pH ~ 4) buffer and monitoring the release of major ions with time 622634635636637. An acidic buffer, to some extent, mimics the acidic environment during osteoclastic activity. In one study, an in vivo behavior of porous β-TCP scaffolds prepared from rod-shaped particles and that prepared from non-rod-shaped particles in the rabbit femur was compared. Although the porosities of both types of β-TCP scaffolds were almost the same, a more active osteogenesis was preserved in the region where rod-shaped ones were implanted 638. Furthermore, the dimensions of both the particles 609 and the surface microstructure 604 were found to influence the osteoinductive potential. These results implied that the microstructure affected the activity of bone cells and subsequent bone replacement.

Fixation of any implants in the body is a complex dynamic process that remodels the interface between the implants and living tissues at all dimensional levels, from the molecular up to the cell and tissue morphology level, and at all time scales, from the first second up to several years after implantation. Immediately following the implantation, a space filled with biological fluids appears next to the implant surface. With time, cells are adsorbed at the implant surface that will give rise to their proliferation and differentiation towards bone cells, followed by revascularisation and eventual gap closing. Ideally, a strong bond is formed between the implants and surrounding tissues 568. An interesting study on the interfacial interactions between calcined HA and substrates has been performed 639, where the interested readers are referred for further details.

The aforementioned paragraph clearly demonstrates an importance of studies on cellular responses to CaPO₄ scaffolds. Therefore, such investigations have been performed extensively for several decades 582640641642643644645646647648649650651652653654655. For example, bioceramic discs made of 7 different types of CaPO₄ (TTCP, HA, carbonate apatite, β-TCP, α-TCP, OCP and DCPD) were incubated in osteoclastic cell cultures for 2 days. In all cases, similar cell morphologies and good cell viability were observed; hoverer, different levels of resorbability of various types of CaPO₄ were detected 643. Similar results were found for fluoridated HA coatings 645. Experiments performed with human osteoblasts revealed that nanostructured bioceramics prepared from nano-sized HA showed significant enhancement in mineralization compared to microstructured HA bioceramics 644. In addition, the influence of lengths and surface areas of rod-shaped HA on cellular response were studied. Again, similar cell morphologies and good cell viability were observed; however, it was concluded that high surface area could increase cell-particle interaction 648. Nevertheless, another study with cellular response to rod-shaped HA bioceramics, revealed that some types of crystals might trigger a severe inflammatory response 651. In addition, CaPO₄-based sealers appeared to show less cytotoxicity and inflammatory mediators compared with other sealers 646. More examples are available in literature.

Cellular biodegradation of CaPO₄ scaffolds is known to depend on its phases. For example, a higher solubility of β-TCP was shown to prevent L-929 fibroblast cell adhesion, thereby leading to damage and rupture of the cells 656. A mouse ectopic model study indicated the maximal bone growth for the 80 : 20 β-TCP : HA biphasic formulations preloaded with human mesenchymal stem cells when compared to other CaPO₄657. The effects of substrate microstructure and crystallinity have been corroborated with an in vivo rabbit femur model, where rod-like crystalline β-TCP was reported to enhance osteogenesis when compared to non-rod like crystalline β-TCP 638. Additionally, using a dog mandibular defect model, a higher bone formation on a scaffold surface coated by nanodimensional HA was observed when compared to that coated by a micro-dimensional HA 658. Furthermore, studies revealed a stronger stress signaling response by osteoblast precursor cells in 3D scaffolds when compared to 2D surfaces 659.

Mesenchymal stem cells are one of the most attractive cellular lines for application as bone grafts 660661. Early investigations by Okumura, et al. indicated an adhesion, proliferation and differentiation, which ultimately became new bone and integrated with HA scaffolds 641. Later, a sustained co-culture of endothelial cells and osteoblasts on HA scaffolds for up to 6 weeks was demonstrated 662. Furthermore, a release of factors by endothelial and osteoblast cells in co-culture supported proliferation and differentiation was suggested to ultimately result in microcapillary-like vessel formation and supported a neo-tissue growth within the scaffold 444. More to the point, investigation of rat calvaria osteoblasts cultured on transparent HA bioceramics, as well as the analysis of osteogenic-induced human bone marrow stromal cells at different time points of culturing indicated to a good cytocompatibility of HA bioceramics and revealed favorable cell proliferation. The positive results for other types of cells have been obtained in other studies 202, 403,404,405, 663,664,665.

Interestingly that HA scaffolds with marrow stromal cells in a perfused environment were reported to result in ~ 85 % increase in mean core strength, a ~ 130 % increase in failure energy and a ~ 355 % increase in post-failure strength. The increase in mineral quantity and promotion of the uniform mineral distribution in that study was suggested to attribute to the perfusion effect 535. Additionally, other investigators indicated to mechanical properties increasing for other CaPO₄ scaffolds after induced osteogenesis 534537.

To finalize this section, one should mention on the recent developments to influence the cellular response. First, to facilitate interactions with cells, the surface of CaPO₄ scaffolds can be functionalized 666667668669670. Second, it appears that crystals of biological apatite of calcified tissues exhibit different orientations depending on the tissue. Namely, in vertebrate bones and tooth enamel surfaces, the respective a, b-planes and c-planes of the apatite crystals are preferentially exposed. Therefore, ideally, this should be taken into account in artificial bone grafts. Recently, a novel process to fabricate dense HA bioceramics with highly preferred orientation to the a,b-plane was developed.

The results revealed that increasing the a,b-plane orientation degree shifted the surface charge from negative to positive and decreased the surface wettability with simultaneous decreasing of cell attachment efficiency

671,672,673. The latter finding resulted in further developments on preparation of oriented CaPO₄ compounds 674675676.