In 2013, Sergey V. Dorozhkin realized an excellent chronological study where he describes the history of CaP since their discovering in bones and teeth until their last applications in medicine [1]. According to this work, in the last quarter of the 17th century appeared the first studies about the structure and composition of bones, teeth and other types of calcified tissues. Must be highlighted the study published by Antonie Philips van Leeuwenhoek in 1677, where he describes the observation of small pipes in the shinbone of a calf, nowadays named Haversian canals because of the British physician Clopton Havers, and transparent globuls in cow teeth or elephants ivories, which was the first recognition of CaP single crystals in bones [1,2]. It was 1769 when the famous Swedish chemist and metallurgist Johan Gottlieb Gahn discovered the first evidence of CaP existence in bones [1]. In 1770, the presence of orthophosphates was also revealed in blood serum [1].

In the 19th century have to be underlined the studies of Fourcroy, who established the chemical composition of teeth with Vauquelin, or Sir Humphry Davy, who established in 1814 the general principles of bone and teeth formation, well-known as biomineralization [1]. He exposed that bones mainly consist on a gelatinous membrane in the earliest period of animal life, which is destined to gradually acquire calcium phosphate giving their subsequent hardness and durability to the bones. Other studies realized during this century conclude interesting observations. For example, CaP is not found in infants [1], or in lower amounts in pregnant women [3] compared to normal adults. Compositional studies in different bones of human body [4] or bones from young and old individuals were also realized [5].

CaP have been widely studied from a biological, structural and morphological points of view [1]. Nowadays, it is considered that CaP are of a special importance since they are the most important inorganic constituents of hard tissues in vertebrates [6]. In the form of a poor crystalline, non-stoichiometric, ion-substituted CDHA (commonly referred to as “biological apatite”), CaP are present in bones, teeth, deer antlers and tendons of mammals to give these organs stability, hardness and function [7].

During last few centuries, CaP have been proposed to treat various diseases such as rachitis scrofula, diarrhea, ulcerations, inflammations, caries or fractures of the bones [1]. Nowadays, they are commonly applied in orthopedic and maxillofacial application as well as in dentistry. Moreover, apart from these applications, they have been recently suggested as vehicles of drugs [8], peptides or ADN molecules [9]. However, it has been the bone tissue engineering, the field in which they are mainly studied to get the ideal implant to regenerate bone tissue while it is resorpted.

The origin of CaP can be natural, for example calcined bovine HA or carbohydroxiapatite from the eggshells [18], and synthetic [19]. Most common synthesis methods are based on precipitation [20] or obtaining from other previous phases by thermal treatment for example. Also can be obtained by less conventional methods such us combustion synthesis [21].

HA, Ca10(PO4)6(OH)2, has been widely used in the field because its chemical composition is quite similar to mineral phase of bones, which can be defined as a CDHA with OH− partially substituted by CO32− groups [22]. HA presents an ionic character and its crystalline structure can be described as a hexagonal packaging of oxygen atoms where metals are placed on the tetrahedral and octahedral sites [19].

HA has been included within the bioactive ceramics along with bioactive glasses and glass–ceramics [19]. However, HA is a very stable phase under the physiological conditions, which is translated into a low solubility (Table 2) and slow resorption kinetics [23]. Most of the bioreabsorbable ceramics, with the exception of plaster or wollastonite, are based on CaP. Their biodegradabily increases following the sequency (HA<<<β-TCP<α-TCP) [19].

In one hand, TCP is the biodegradable CaP par excellence [19]. It belongs to the Whitlockites family which correspond to the general formula (Ca,Mg)3(PO4)2. It presents three polymorphs: β, α and α′, from the lowest to the highest stability temperatures and being the last one total reversible in both senses [19]. In case of α and β, pure phases as well as controlled mixtures may be obtained with adequate thermal treatment and amount of Mg [24]. β→α-Ca3(PO4)2 transition temperature is increased with Mg (Mg3(PO4)2) and Mg and Si (CaMg(SiO3)2) content [25]. Moreover, β-TCP phase can be also stabilized with partial Ca substitution for Sr and Zn and α-TCP with partial PO43− substitution for SiO44−[26].

Wherever α or β phase, possess higher solubility than HA (HA<β-TCP<α-TCP) as shown in Table 2 and according with the literature [27]. Therefore, the biphasic calcium phosphates (BCP) concept appear to determine the optimum balance of a more stable phase and a more soluble phase with the aim to control in vivo bioresorbability through the phase composition [28]. The vast majority of CaP bioceramics are based on HA [29] both types of TCP [26] and multiphase formulations thereof [12,30].

As other formulations, CaP might be processed in both dense and porous forms in bulk, as well as in the forms of powders, granules, scaffolds or coatings [12]. Depending of the requirements they have to be processed in different forms. For example, improved mechanical properties require dense materials while cell colonization, bone regeneration and vascularization require 3D porous structures such as scaffolds [42] or granules [43]. Powders can be used, for example as raw materials to obtain self-setting formulations. Nanoparticles or nanostructured particles can be used not only to obtain dense materials [44] but also for drug delivery systems or gene transfection [8,9].

Processing normally consist in two steps, conformation and following sintering. Most common conformation methods include: uniaxial compaction [45], isostatic pressing [46], granulation [47], slip casting [48], gel casting [49], freeze casting [50], atomization [51], polymer replication [52], extrusion [53], low pressure injection [20], slurry, dipping or thermal spraying [54]. Also, shaping can be realized during cooling and solidification of previously melted raw materials [12]. It has to be noted that forming and shaping of any ceramic products require a proper selection of the raw materials in terms of particle sizes distribution [12].

Sintering is a processing step with great interest to consolidate the shaped forms. During the sintering of CaP, different processes take place [12]: humidity, carbonates and remaining volatile chemicals from synthesis are removed; shrinkage occurs as a consequence of the removal of these gases while powders densify, and grain size increases. Chemical changes and chemical decomposition of all acidic orthophosphates by water lost take place [12].

To obtain improved mechanical properties grain growth and shrinkage must be minimized during the sintering. For this reason, sintering methods such as hot isostatic pressing [46], microwave [55], spark plasma sintering [56] or rate controlled sintering have been suggested.

Nowadays, implants with even more complex geometries can be developed thanks to the computer-aided design and manufacturing [12]. For example, an image of a bone defect in a patient can be used to develop a three-dimensional (3D) structure [57]. The computer reduces the model to slices or layers and 3D structures are constructed layer-by-layer. Laser sintering [58], laser cladding [59], 3D printing [60], solid freeform fabrication [61] and stereo lithography [62] are some examples of this kind of technology.

Finally, polymeric–ceramic composites have increased the processing possibilities for ceramics. These kinds of hybrids have been widely studied and developed for biomedical applications. The main reasons to add a polymeric phase are: the improvement of the mechanical properties regarding to the brittleness of ceramics, their easy fabrication as 3D structures with the possibility to include drugs or molecules in their matrix and their ability to control their delivering rate. Ceramic powders, granulates or 3D structures can be covered or embed in a polymer matrix. Polymeric processing methods such as freeze–thawing [63], photopolymerization [64], melt-extrusion [65], have been used in the processing of polymeric–ceramic composites.

Bioceramics can be divided in the following groups according with their behavior once they are implanted in the living organisms [19]: Bioinert: They do not show important changes and neither interaction with living tissue; Biotoxic: They release substances in toxic concentrations and/or trigger the formation of antigens that may cause immune reactions; Biocompatible: They do not provoke adverse reactions and neither release any toxic constituents; Bioactive: They promote formation of a surface layer of biological apatite responsible of bone-bonding; Biodegradable or bioresorbable: They degrade with time in presence of physiological fluids.

The biological response of implanted CaP follows a similar mechanism than fracture healing. It includes a hematoma formation, inflammation, neovascularization, osteoclastic resorption and a new bone formation [12]. In general, CaP show different bioactivity and bioresorbability behaviors and depend on the Ca/P ratios [19], porosity grade or densification [87] and solubility (Table 2).

Bioactivity

When bioactive materials are implanted interact with physiological media and surrounding bone tissue. A CDHA layer precipitates at the bone tissue–biomaterial interface and it is responsible of the bone bonding. Crystal phases, surface roughness and porosity are parameters that play an important role in the bioactive behavior [12].

Because of the great importance of the bone bonding ability of CaP, the study of the CDHA layer at the bone tissue–biomaterial interface and its formation have focused an extraordinary attention of researchers. Several in vitro and in vivo studies have been carried out to study the bioactivity of different materials. For in vitro studies, different type media, which pretend to simulate body fluids, have been developed. Ringer solution or Tris–HCl are some examples. A well known procedure was the developed by Kokubo. The first version of simulated body fluid (SBF) arrived in 1990 [92]. Based on Kokubo studies and with the aim to perform comparative studies in the international setting, it was published in 2007 the ISO/FDIS 23317 normative: Implants for surgery—In vitro evaluation for apatite-forming ability of implant materials. After in vitro studies, the most common method to observe new formed CDHA is SEM.

The process has been widely studied and described in detail for bioactive glasses [93]. Bioactivity has been also studied for CaP ceramics and some mechanisms has been suggested [94,95]. For example, in vitro biodegradation and bioactivity of a β-Ca3(PO4)2–CaMg(SiO3)2 ceramic with the eutectic composition (Fig. 2) has been studied [96]. Also, the bioactivity of a glass-ceramic (Fig. 3) and a ceramic with the same 45% CaO, 22% SiO2, 28% P2O5 and 5% MgO (wt.%). based composition, selected in the Ca3(PO4)2–CaSiO3–3MgO·4SiO2 pseudo-ternary system, was studied and compared in SBF [97].

Fig. 2.

Low magnification FE-SEM microphotograph of the polished and thermically etched surface (1100°C) of 60wt% Ca3(PO4)2–40wt% CaMg(SiO3)2 composition sintered 4h at 1225°C. TCP correspond to tricalcium phosphate phase and D to the diopside [96].

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Fig. 3.

HA layer precipitated on the surface of a glass ceramic material based on TCP after in vitro assay on SBF [97]. (A) HA layer is cracked and partially removed, showing the glass over which has precipitated, (B) detailed crack of HA layer and (C) detailed crystals of HA layer.

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Development of nanomaterials, their singular properties and sophistication have revolutionized numerous research fields. They can develop more complex functions and be designed to mimic the physicochemical properties of natural materials [126].

Conventional synthesis routes obtain materials on the scale of millimeters or larger and then milled to have micrometer or nanometer scale features. However there are bottom-up approaches which are inspired on the nature. Natural materials such as proteins, DNA or RNA, are built by self-assembly [130]. Most of the different methodologies proposed to prepare nanodimensional and/or nanocrystalline apatites are inspired on the bottom-up approach [126]. There is a wide range of bottom-up methods. The wet chemical precipitation is the most used [131]. Morphology, phase formation and crystallinity of precipitated apatites have been studied in relation with temperature [132], calcium ion concentration [133] or pH [134]. Other method commonly used in the synthesis of nanosize CaP is the hydrothermal [135]. Fine-grained single crystals, characterized by high purity, controlled morphology and narrow size distribution can be obtained [136]. Other alternative synthesis methods can be found in the literature to prepare nanodimensional CaP. Some examples are sol–gel [137], solution combustion method (Fig. 4) [21], mechanochemical [138] or microwave are some examples [139].

Fig. 4.

Synthesis of HA by combustion in (A) aqueous media and (B) oxidizing media [21].

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Moreover, methods based on the natural nucleation and growth of nanocrystalline apatite between self-assembled collagen fibrils, have been also developed. This process has been simulated using macromolecules as templating agents. Precipitation of nanodimensional apatites from aqueous solutions has been carried out, for example, in the presence of dissolved polyacrylic acid [140], polyvinyl alcohol [141], amino acids [142], or gelatin [143]. The aim of this type of synthesis is the obtention of biocomposites with structure and mechanical properties similar to bones. Moreover, it allows the obtention of complex morphologies [126]. An interesting example is a pH-induced self-assembly peptide-amphiphile, to obtain nanostructured fibrous scaffolds by cross-linking, and following mineralization. C-axes of CDHA crystals grow aligned with the long axes of the fibers similar to biological apatite in bones [144].

According with the properties previously described, applications of CaP nanomaterials, are focused in two main directions: in biocomposites and hybrid to enhance bioactivity [150]; and in the manufacturing of dense compacts or porous scaffolds with improved mechanical properties [86].

In dentistry, CaP and nanodimensional apatite within them, have been suggested for enamel remineralization [151]. Enamel is mainly composed by inorganic nanodimensional apatite and difficult to be self-repaired by living organisms after bacteria damage. Most conventional treatments are based on materials fillings. However, frequently secondary caries arise at the tooth–materials interface [152]. Nanodimensional HA and CDHA have been encouraging proposed due to the chemical and phase similarities [112].

In other hand, CaP have been used as vehicle of various agents with drug delivery, gene therapy or bioimaging diagnosis purposes. CaP provide to these drugs or genes a protective environment that avoids their degradation, make easier their penetration trough the cell membrane and control the release [153]. Added to the biocompatibility and biodegradability nature of CaP, nanoparticles possess a huge specific surface area and small particle size. Due to the large specific surface area, nanodimensional CaP can load high amounts of drugs or genes while their small particle size increases the penetration rate. Moreover, after transfection, these particles dissociate into calcium and orthophosphate ions which are found in every cell. For these reasons, the use of CaP nanoparticles for transfection of DNA or RNA into nuclei of living cells, with the aim of repairing missing cell function by enhancing or silencing gene expression, is becoming soon a therapeutic reality. The strategies employed to load different molecules, drugs, genes or radioisotopes divided into two main lines: in situ loading, which consists on load these agents during the synthesis; or load the agent after synthesis [152,154].

Self-setting CaP formulations are mixtures of CaP powders and aqueous solutions such as distilled water [89], SBF [156], aqueous solutions of sodium orthophosphates [157] or H3PO4[158] or citric acid [159].

Self-setting process goes through two possible reactions: acidic-base o hydrolysis. In the first case, CaP with relatively acidic behavior reacts with relatively basic one to produce a relatively neutral compound [89,160]. The second type consists on the hydrolysis of metastable CaP in aqueous media [89,161].

In both cases, when CaP are placed in the aqueous environment they dissolve and mass transport occurs until supersaturated microenvironment is created with regards to the final product [162]. Crystals obtained by precipitation grow and form microneedles and micropellets of CDHA and CDPD, which provide rigidity to the hardened material [89]. Depending of the product obtained after setting, cements are divided in two main types: apatite forming formulations and brushite forming formulations.

In case of apatite-forming formulations, the products of the setting reaction are poorly crystalline HA, CDHA or carbonate apatite, if the setting reaction occurs in presence of carbonates [163]. It is thought that the similar chemistry of those products and biological apatite is responsible of their excellent in vivo resorption [89]. In addition, the setting process occurs at pH values close to the physiological and it contributes to their good biocompatibility [164].

Solubility of these formulations in aqueous solutions is expected to be similar to that of bone mineral. Solubility increases when pH is slightly acid. Moreover, controlled dissolution can be aided by osteoclast [165]. Moreover, crystals obtained can be easily detached, especially after dissolution and it makes the ingestion by osteoclast and other cells easier [166].

Regarding to the setting time, which is stretched on in vivo conditions, it is quite long for apatite forming formulations [89]. It can entail technical complications and, for this reason, an interesting goal consists on decrease their setting time. Different approaches have been suggested, for example, by reducing the amount of liquid phase as low as possible or using additives like H3PO4 to decrease the pH and promote the dissolution of CaP [167].

In this case of brushite forming formulations, the major product of the setting reaction is DCPD [89]. These setting reactions occur by acidic-base process. DCPD only precipitates under pH 6. Consequently, the reaction environments in these formulations are always slightly acidics [158].

Apart from its biocompatibility [89], Brushite presents higher solubility compared with CDHA [168]. It is translated in a faster degradation and quickly resorption in vivo[169]. The disadvantage is the short setting time, which complicates the injectability. For this reason, one of the main goals consists on decrease the viscosity [159].

Finally, monetite based formulations have been developed by acidic-base process. Excess of HA is mixed with H3PO4. The aim is obtain a phase mixture to control the degradability rate. They pretend to be applied as granulates which provide a 3D structure [85].

In this section, the properties of CaP self-setting formulations will be described regarding to their biomedical and clinical applications. They should be easily injected and once they are implanted, setting and hardening processes must occur in reasonable times. They must possess certain mechanical resistance to avoid mechanical failures and they have to be reabsorpted while new bone is formed.

Combination of self-setting nature with their high biocompatibility, bioresorbability as well as osteoconductivity previously described make CaP cements encouraging for bone augmentation and bone defect healing applications [89]. Self-setting CaP have been successfully used for different fracture treatments [188] augmentation of osteoporotic vertebral bodies [89] as well as vertebroplasty and kyphoplasty for osteoporosis-induced vertebral compression fractures [189]. In addition, they are used to aid the fixation of titanium implants and screws in mechanically poor bone [190] as well as to decrease pains associated with unstable vertebrae. Furthermore, they are used for bone augmentation and bone defect healing in oral, maxillofacial and craniofacial surgery [89]. It has to be noted that these applications do not require materials with mechanical or stress resistance.

Other important application for CaP cements is as drug delivery system. They add attractive properties such as injectability, biodegradability and large surface area [191] to the general requirements of a drug carrier which must incorporate, retain and gradually deliver certain drugs or molecules (Fig. 5). For these reasons, self-setting CaP are encouraging candidates as drug carriers and they have been widely studied with this purpose. There are three main research lines [89] (Table 4).

Fig. 5.

Different mechanism of drug release using CaP cements as vehicles. (A) Diffusion mechanism, (B) diffusion mechanism where the diffusion rate is slower due to the HA layer formed as consequence of bioactivity and (C) release due to biodegradation process of the particle.

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Self-setting CaP formulations are promising materials used as drug-delivery systems of antibiotics, anticancer, anti-inflammatory agents or bone morphogenetic proteins in the treatment of various bone diseases, such as tumors, osteoporosis and osteomyelitis [192,193]. Also CaP cements with Zn contain, have been designed to deliver this Zn with antibacterial purposes [194].

Inside of this new research area, the surface engineering, several methods have been developed to obtain the desirable properties in the surface of materials maintaining their bulk. They can be classified into two different types: coating techniques and conversion or modification techniques [119]. The first one consists in a coating over the surface of the substrate with new material. Coating technology is applied to cover metallic implants with CaP. The latter one, consists on a conversion or modification of the material surface [196].

A coating can be prepared as monolayer or multilayer coatings which are produced by a layer-by-layer deposition. Multilayer structures provide different or graded compositions and/or properties [119]. Regarding the thickness there are thin layers, ranging from nanometer to several micrometers thickness, and a thick layer, with thickness exceeding several micrometers [119].

Ceramics and glass ceramics CaP based compositions have been used to prepare coatings [197]. The most important coating techniques are classified in Table 5[119,198].

CaP coatings have been mainly thought to work on the bone–metallic implant interfaces. Once deposition is realized, the requirements are a suitable adhesion and cohesion between the coating and the metallic implant. Moreover, thickness will play an important role regarding the mechanical properties, the bioactivity and the biodegradation.

Between the different applications of CaP in the biomaterials field must be highlighted its contribution as bioactive coatings. It means that their function consists in promoting bone bonding to ensure a properly fixation of implants. In a first moment, CaP are suggested as coatings for metallic implants, which are directed to load-bearing applications. Moreover, CaP coatings reduce the ion release from metallic surface [206], slow down metal surface degradation [207] and corrosion and enhance biocompatibility [208].

Between CaP, HA is the bioactive material most commonly used. Also, bioactive glasses with CaO and P2O5 contain has been suggested for bone adhesion [209]. However, they show a quickly dissolution which can cause the instability of the implant with time [10].

Multilayer coatings are applied to promote a quickly reactivity of the surface in the first moments, which ensure bone-bonding and fixation in short periods decreasing the immobilization time, while the bottom layer ensure the stability of the implant. For example, a double layer of HA (bottom layer) and HA/Glass (surface layer) have been developed with this propose [10].

Bioactivity is the most characteristic property of these glass and glass–ceramic compositions, as their name indicates, and the main reason why these compositions are studied and developed. Bioactive glass and glass–ceramics clearly show formation of a HA layer on their surfaces, when they are immersed in simulated body fluid (SBF) [211]. Moreover, in vivo studies have shown that bioactive glasses bond to bone more rapidly that other bioactive ceramics [212].

An interesting study was realized to compare the bioactive behavior of a polycrystalline and glass-ceramic material with same oxides compositions. The composition of both materials was based on a 38% CaO, 27% P2O5, 24% SiO2 and 9% MgO (wt%) [97]. While polycrystalline material, obtained after thermal treatment at 1050°C, presents a composition mainly based on a 60% of β-TCP (where Ca2+ is partially substituted by Mg2+) and Wollastonite (W) as secondary phase, the glass-ceramic, obtained by quenching after heating at 1600°C, presents a mixture of crystalline phases: β-TCP, with Mg2+ solid solution, and α-TCP embed in a glass matrix of alkaline earth silico phosphates. Glass–ceramic shown higher bioactivity, in terms of HA layer formed, than the polycrystalline material. It was associated to the more amorphous structure, glass matrix of alkaline earth silico phosphates, which is easier attacked.

Apart from their excellent bioactivity, their solubility products possess osteogenic properties [212]. However, their bioresorption is slow and the bulk of these materials remain unchanged. For this reason, an appropriate porous structure is recommended to improve the ingrowth of new bone into the implant and its resorption [118]. However, the production of 3D porous structures involves a sintering step and sintering can comprises phase crystallizations. [212] The overlap of the sintering and crystallization events must be avoided while the densification process occurs. Sintering should precede crystallization in order to obtain mechanically strong amorphous or partially crystallized scaffolds. Alkali-free bioactive diopside–tricalcium phosphate glass–ceramics in the system CaO–MgO–SiO2–P2O5 have been studied for scaffold fabrication. Good densification was achieved for glass powder compacts obtaining sintered and mechanically strong glass–ceramics. Also compositions with MgO partially substituted by ZnO have been developed.

Otherwise, process developments in foaming, solid freeform fabrication and nanofiber-spinning have now allowed the production of porous bioactive glass scaffolds from both melt- and sol–gel derived glasses [212].

Another alternative to overcome this problem consists on develop an in situ porous structure when they are implanted [118]. To fulfill this approach, at least two phases are necessary, one bioactive and the other resorbable [118]. At the same time the microstructure should be as homogeneous as possible and easily controlled [118]. These features can be achieved in binary eutectic structures. For example, the binary system wollastonite (W)-tricalcium phosphate (TCP) was selected [213], where W is bioactive [214,215] and TCP is resorbable [210] and presents eutectic point at 1402±3°C for the composition 60wt% W and 40wt% TCP [213]. To obtain a eutectic structure, heat extraction must take place from the eutectic liquid composition. De Aza et al. selected radial heat flux in order to achieve microstructures formed by rounded colonies named rosettes (Fig. 6) [118].

Fig. 6.

SEM micrographs of bioeutectic after etching in diluted acetic acid. Micrographs show microstructures formed by rounded colonies of porosity named rosettes [216].

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When the eutectic composition is immersed in SBF, three events occur on the periphery of the sample [118]: the dissolution of the wollastonite, the pseudomorphic transformation of the TCP to HA and the superficial deposition of HA. At the beginning, two quasi-simultaneous mechanisms take place. First, W phase in contact with the SBF exchange two H+ from the SBF for one Ca2+ from the W network. W crystals are transformed into an amorphous silica phase [215,217]. This reaction increases the pH at the material-SBF interface and Ca2+ and silicon ions are released into the SBF media. Simultaneously, the TCP starts to react with the Ca2+ and OH− ions present in the medium. It gives rise to the pseudomorphic transformation of TCP into HA. The remaining silicon and Ca2+ ions are caught by the phosphate at the material-SBF interface. Initially, the precipitation occurs on the neoformed HA lamellae, which act as nuclei. Later, this precipitation is increased covering the surface of the sample [118].

A disadvantage of bioeutectis is the pore sizes created in situ by biodegradation. These porosity sizes are around 1μm width and 10μm length. Cells cannot colonize the material with this pore sizes. However, it is enough to aids bioresorption. Bioetectics are promising materials and show encouraging results regarding bone-like forming and human bone marrow cells proliferation and growth. Deeper investigations have been made around the improvement of bioactivity and mechanical properties [218,219]. Bioactivity, biodegradation, porosity as well as mechanical properties are directly related between then and dependent from the microstructure [220,221]. This microstructure can be controlled but well-founded knowledge of nucleation and growth mechanism of the crystal phases is required. For this reason, devitrification and crystallization study of W-TCP eutectic glass was realized [218]. Also Mg was introduced in order to reduce glass network, whit the aim to increase bioactivity and biodegradability and improve mechanical properties [219].

Finally, alkali-free bioactive diopside–tricalcium phosphate glass–ceramics in the system CaO–MgO–SiO2–P2O5 have been studied for scaffold fabrication [222]. This study is based on the idea that the overlap of the sintering and crystallization events must be avoided while the densification process occurs. Sintering should precede crystallization in order to obtain mechanically strong amorphous or partially crystallized scaffolds. Good densification was achieved for glass powder compacts obtaining well-sintered and mechanically strong glass–ceramics. Also compositions with MgO partially substituted by ZnO have been developed [223].

Bioglasses and glass–ceramics are mainly used as dense materials or coatings. Regarding dense bioglasses and glass–ceramics have been used for dental implants, maxillofacial reconstruction, vertebrae union or middle ear reconstruction. As coatings, they have been mainly used to fix metallic implants and prosthesis to bones, for example, hip prosthesis.