Ionic substitutions in calcium phosphates synthesized at low temperature

Acta Biomaterialia 6 (2010) 1882–1894

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Review

Ionic substitutions in calcium phosphates synthesized at low temperature E. Boanini a, M. Gazzano b, A. Bigi a,* a b

Department of Chemistry ‘‘G. Ciamician”, University of Bologna, 40126 Bologna, Italy ISOF-CNR, c/o Department of Chemistry ‘‘G. Ciamician”, Bologna, Italy

a r t i c l e

i n f o

Article history: Received 5 November 2009 Received in revised form 16 December 2009 Accepted 21 December 2009 Available online 28 December 2009 Keywords: Calcium phosphates Ionic substitutions Hydroxyapatite Bone cements Biomimetic coatings

a b s t r a c t Ionic substitutions have been proposed as a tool to improve the biological performance of calcium phosphate based materials. This review provides an overview of the recent results achieved on ion-substituted calcium phosphates prepared at low temperature, i.e. by direct synthesis in aqueous medium or through hydrolysis of more soluble calcium phosphates. Particular attention is focused on several ions, including Si, Sr, Mg, Zn and Mn, which are attracting increasing interest for their possible biological role, and on the recent trends and developments in the applications of ion-substituted calcium phosphates in the biomedical field. Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction Calcium orthophosphates (CaPs) are widespread in nature; in particular, they constitute the main inorganic components of the hard tissues of vertebrates. The mineral phase of bone and teeth is a basic calcium phosphate, which is assimilated to synthetic hydroxyapatite (HA, Ca10(PO4)6(OH)2) [1–3]. However, biological apatites differ from stoichiometric HA in several respects, including non-stoichiometry, small crystal dimensions and poor crystallinity (i.e. low degree of structural order). Biological apatites are indeed carbonated apatites, and they contain significant amounts of further foreign ions, either incorporated in the apatite crystal lattice or just adsorbed on the crystal surface. The mineral phases of bone, enamel and dentin exhibit slightly different ionic composition (Table 1), crystallinity and solubility [8]. The low degree of crystallinity and the small crystal sizes—typical dimensions of bone crystals are 50  25  4 nm [9]—are important factors related to the relatively high solubility of biological apatites with respect to stoichiometric HA. Further calcium phosphates occurring in biological tissues, usually in pathological calcifications, include amorphous calcium phosphate (ACP), dicalcium phosphate dihydrate (DCPD), octacalcium phosphate (OCP), Mg-substituted b-tricalcium phosphate (b-TCMP) (Table 2). Thanks to their similarity with the inorganic phases of biomineralized tissues, CaPs display excellent biocompatibility and bioactivity, which justify their widescale use in the preparation of biomaterials for hard tissue substitution and repair. Interest is ad* Corresponding author. Tel.: +39 051 2099551; fax: +39 051 2099456. E-mail address: [email protected] (A. Bigi).

dressed not only towards HA, but includes further more resorbable calcium phosphates, such as ACP, a-TCP, b-TCP, OCP and DCPD, that can easily transform into nanocrystalline apatite [10,12–14]. Nucleation and growth of CaPs in biological systems occur in an environment rich in ions, which can affect both the kinetics and the thermodynamics of crystallization, and then the relative stability of CaPs [15]. It follows that the study of ionic substitution in CaPs is relevant for several reasons, including a better understanding of the biomineralization processes, control of the properties of the precipitated phase, increase of the bioactivity of the material, and delivery of ions able to treat bone diseases. This review is focused on ionic substitutions in CaPs prepared at low temperature, i.e. synthesized in aqueous medium or prepared through hydrolysis of more soluble calcium phosphates. First, the main characteristics of CaPs, and the main modifications induced by ionic substitutions are briefly reviewed. In the next section, further details are reported for some selected ions, namely Sr, Mg, Si, Zn and Mn, which have recently attracted remarkable interest for their possible biological relevance. The final section reports the results achieved in a few biomedical applications where the effect of ionic substitutions has been tested. 2. Calcium orthophosphates 2.1. Hydroxyapatite HA (Ca10(PO4)6(OH)2) is the most stable, the most dense and the most insoluble of the CaPs. From a chemical and structural point of view it is the material most similar to the mineral component of bones and teeth. HA can be easily prepared in aqueous solution.

1742-7061/$ - see front matter Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2009.12.041

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E. Boanini et al. / Acta Biomaterialia 6 (2010) 1882–1894 Table 1 Comparative composition of human enamel, dentin and bone.

a

Ca (wt.%) P (wt.%)a CO2 (wt.%)a Na (wt.%)a K (wt.%)a Mg (wt.%)a Sr (wt.%)a Cl (wt.%)a F (wt.%)a Zn (ppm)b Ba (ppm)b Fe (ppm)b Al (ppm)b Ag (ppm)b Cr (ppm)b Co (ppm)b Sb (ppm)b Mn (ppm)b Au (ppm)b Br (ppm)b Si (ppm) Ca/Pa

Enamel

Dentine

Bone

37.6 18.3 3.0 0.70 0.05 0.2 0.03 0.4 0.01 263 125 118 86 0.6 1 0.1 1 0.6 0.1 34

40.3 18.6 4.8 0.1 0.07 1.1 0.04 0.27 0.07 173 129 93 69 2 2 1 0.7 0.6 0.07 114

36.6 17.1 4.8 1.0 0.07 0.6 0.05 0.1 0.1 39c

1.59

1.67

0.33d <0.025d 0.17d

500e 1.65

a Data in these rows are from Ref. [3]. The composition is relative to the ash content of whole enamel (96%) and of the whole dentin (73 wt.%). Bone is from bovine cortical bone. The percentage is relative to the mineral fraction (73 wt.% of dry fat-free bone). b Data in these rows are from Ref. [4] unless otherwise indicated. The composition is relative to human teeth. c From Ref. [5]. d From Ref. [6]. The composition is relative to wet human bone. e From Ref. [7]. The value is relative to the mineralizing osteoid regions of normal tibiae from young mice and rats.

The literature reports a variety of syntheses that, although differing in terms of the detailed conditions, essentially follow one of two methods. The first one consists in the stoichiometric titration of a calcium hydroxide slurry with phosphoric acid up to neutrality [16]. In the second one, the precipitation method, a solution of ammonium or sodium monohydrogen phosphate is added dropwise to a calcium solution (generally calcium nitrate or acetate) or vice versa [17,18]. Both methods yield good crystalline products when carried out at 100 °C. Lowering the temperature provokes an increase in the amount of amorphous material and/or decrease in crystallinity. The titration method is performed in a basic environment, whereas the precipitation method often implies addition of ammonia or sodium hydroxide to maintain a basic pH (>9) during precipitation. A further factor that influences crystallinity and crystal dimensions is the reaction time: short reaction times provide nanocrystals and/or amorphous products, whereas crystal dimensions and crystallinity increase with increasing reaction time. Syntheses in air usually mean incorporation of carbonate ions in the reaction products. Apatites, as well as other CaPs, can also be obtained as the result of the transformation of more soluble, metastable phosphates in a wet environment. The main methods of preparation of CaPs are summarized in Table 2. Of the two known crystal forms of HA—monoclinic, space group P21/b, and hexagonal, space group P63/m—only the hexagonal phase is of practical importance because the monoclinic form is destabilized by the presence of even small amounts of foreign ions [3]. The hexagonal structure contains two different cation sites, Ca(I) and Ca(II), but only one phosphate environment (Ca(I), Ca(II) are used for stoichiometric apatite; M(I), M(II) are the general sym-

Table 2 Properties, preparations and occurrence of the biologically relevant phosphates. Abbreviation Formula

a

c

Ca/P ratio pKsp (25 °C)a

pH stabilitya

Main preparation methods

Hydroxyapatite

1.67

116.8

9.5–12

(i) Titration of Ca(OH)2 with H3PO4

Occurrence in biological tissues

HA

Ca10(PO4)6(OH)2

OCP

Ca8H2(PO4)6
5H2O Octacalcium phosphate

1.33

96.6

5.5–7.0

Dropwise addition of Ca(Ac)2 to

b-TCP

Ca3(PO4)2

b-Tricalcium phosphate (whitlokite)

1.5

28.9

b

(i) Solid-state reaction of CaCO3 and DCPD at 900 °C (ii) Thermal conversion of CDHA

Dental and urinary calculi, soft-tissue deposits, artritic cartilage, usually present as b-TCMP

a-TCP

Ca3(PO4)2

a-Tricalcium phosphate

1.5

25.5

b

Heat treatment of b-TCP at 1300 °C

Not found

Amorphous calcium phosphate

1.2–2.2

c

b

Fast mixing of Ca2+ and

Soft-tissue calcifications

ACP

b

Name (mineral)

Cax(PO4)y
nH2O

Bone, dentin, enamel, dental calcifications, (ii) Dropwise addition of HPO24 urinary stones, solution to Ca2+ solution, pH > 9 (iii) Hydrolysis from other phosphates atherosclerotic plaques Dental and HPO24 /H2 PO4 solutions at 60 °C, pH 5 urinary calculi

HPO24 solutions, RT

MCPM

Ca(H2PO4)2
H2O

Monocalcium phosphate monohydrate

0.5

1.14

0–2

Titration of H3PO4 with Ca(OH)2 in strong acidic environment

MCPA

Ca(H2PO4)2

Anhydrous monocalcium phosphate

0.5

1.14

b

Heat treatment of MCPM at T > 100 °C Not found

DCPD

CaHPO4
2H2O

Dicalcium phosphate dihydrate (brushite)

1.0

6.59

2–6

Dropwise addition of a Ca2+ solution Dental calculi, urinary stones to a HPO24 solution at 60 °C, pH 4 chondrocalcinosis

DCPA

CaHPO4

Anhydrous dicalcium phosphate (monetite)

1.0

6.90

b

Heat treatment of DCPD at T > 100 °C Not found

TTCP

Ca4(PO4)2

Tetracalcium phosphate

2.0

38–44

b

Solid-state reaction of DCPA with CaCO3 at high T

Not found

CDHA

Ca10 x(HPO4)x (PO4)6 x(OH)2

Calcium-deficient hydroxyapatite

1.5

85.1

6.5–9.5

Hydrolysis of ACP or a-TCP

Not found

x

Data from Ref. [10]. Phases obtained by solid-state reaction or heat treatment of other phases. Cannot be measured precisely, however the following values were reported: 25.7 (pH 7.40), 29.9 (pH 6.00), 32.7 (pH 5.28) [11].

Not found

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bols for substituted apatites). The structure can be roughly described as a phosphate assembly crossed by parallel channels filled by OH ions and parallel to the crystallographic c-axis (Fig. 1). The channel walls are formed of Ca(II) atoms arranged in staggered triangular arrays. Ca(I) atoms have a different environment and are positioned in columns parallel to the OH channels [19]. A unit cell accommodates a formula unit Ca10(PO4)6(OH)2. Among the 10 cations, the 4 Ca(I)s are tightly bonded to 6 oxygens and less strongly to the other 3 oxygens (mean Ca(I)–O distance 0.255 nm), whereas the 6 Ca(II) atoms are surrounded by 7 oxygens (mean Ca(II)–O distance 0.245 nm). Ca(I) atoms are strictly aligned in columns and any small change in the metal–oxygen interactions affects the entire lattice. However, the Ca(II) atoms belonging to consecutive layers are staggered, allowing random local misplacements without compromising the whole structure. As a consequence, cations smaller than Ca or also low concentrations of slightly larger cations are preferably accommodated in site Ca(I) where stronger interactions are present, while larger cations should be accommodated in position Ca(II), even at high concentrations. 2.2. Octacalcium phosphate At physiological pH, OCP (Ca8(HPO4)2(PO4)4
5H2O) is more soluble than apatite, but less than the other CaPs. Its triclinic struc accommodates 8 different Ca environments ture, space group P 1, and 6 phosphate groups. Six of the Ca2+ ions and two of the phosphate groups are hosted in a layer (apatitic layer) where they occupy almost the same positions as in HA structure, as shown in Fig. 2. The apatitic layers alternate with the hydrated layers, which contain the other two Ca2+ ions and the other phosphate group, as well as H2O molecules [20]. Due to its structural features, OCP is often found as an intermediate phase during the precipitation of the thermodynamically more stable HA [3]. It has been detected in a crystal of calcifying dentin and in several pathological calcifications, as urinary and dental calculi [21–24]. Moreover, OCP is believed to play an active role in mitral valve calcification [25]. 2.3. Tricalcium phosphates At least two different crystalline TCPs of stoichiometric composition (Ca3(PO4)2) are known at ambient condition. The rhombohedral b-TCP, space group R3c, can be obtained only at high temperature, by solid-state reaction or by thermal conversion of other phosphates, and is stable up to 1125 °C. It has been identified

Fig. 2. A view of the OCP structure down the c-axis. The hydrated and apatitic layer are highlighted. The positions of Ca atoms connected by black lines and of the phosphate groups of the apatitic layer are very close to those found in the HA structure.

in several pathological calcifications, usually with a relatively high degree of Mg substitution for Ca. In that case the name of the corresponding mineral whitlockite, or the abbreviation b-TCMP, is often preferred. Above 1125 °C b-TCP converts into a-TCP [3]. The unit cell of a-TCP is monoclinic, space group P21/a. The morphology of a-TCP particles is typical of solid-state reaction products, and is quite different from that of CaP crystals synthesized at low temperature, as shown in the comparison with HA, OCP and DCPD crystals reported in Fig. 3a–d. a-TCP is metastable at room temperature even if the transformation does not occur in measurable times in dry conditions. The solubility of both phases is intermediate between orthophosphates, although a-TCP is much more reactive in aqueous solutions than b-TCP and easily hydrolyzes to more stable CaPs. This feature is exploited in the preparation of a-TCPbased cements. 2.4. Amorphous tricalcium phosphate The stoichiometry of TCP can be found also in ACP. However, the abbreviation ACP is generally used to indicate a phase of composition Cax(PO4)y
nH2O with uncertain non-crystalline structure. This usually is the phase that immediately forms in supersaturated solutions obtained by fast mixing of Ca and phosphate ions or in some biological calcifications. In spite of numerous studies, its characterization is still a much-discussed question [26]. It was proposed that the basic structural units of ACP are roughly spherical clusters 0.95 nm in diameter, randomly packed in spherical particles with water in the interstices [27]. Generally accepted points are: (i) an amorphous X-ray diffraction (XRD) pattern but an apatitic-like short-range order; (ii) an usual spherical morphology (diameters in the range 20–200 nm); and (iii) composition depending on the pH and concentrations of Ca and PO34 ions in the mother liquor. The variable composition influences the Ca/P ratio, which very often is around 1.5, though values of 1.2 as well as >2 have also been reported. ACP is highly unstable and hydrolyzes almost instantaneously to more stable CaPs, although it may persist for appreciable periods of time under specific experimental conditions [28]. 2.5. Monocalcium phosphates

Fig. 1. A view of the HA structure along the c-axis. Black lines connect Ca(I) columns in hexagonal networks. Cyan and magenta triangles connect staggered Ca(II) atoms lying in the same plane, but at different height with respect to the c-axis.

Monocalcium phosphate monohydrate (MCPM, Ca(H2PO4)2
 space group. MCPM is the H2O) crystallizes in the monoclinic P 1 most soluble and most acidic phosphate; it converts in the anhydrous form (Ca(H2PO4)2, MCPA) when treated at over 100 °C. MCPM is never found in biological samples. Anhydrous monocal-

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Fig. 3. (a) TEM image of HA nanocrystals; SEM images of (b) OCP crystals; (c) DCPD crystals; (d) a-TCP particles.

cium phosphate (MCPA) has a monoclinic structure, space group  Like the hydrated form, MCPA does not occur in calcified bioP 1. logical tissues. It is not employed in biomedical applications.

of its strong basicity, TTCP reacts with water to yield HA and calcium hydroxide. As a consequence, it is not found in biological calcifications, but it is widely employed for the preparation of TTCPbased bone cements.

2.6. Dicalcium phosphates DCPD (CaHPO4
2H2O) is also known by its mineral name, brushite. DCPD crystallizes in the monoclinic Ia space group. It is the predominant phase below pH 6.5, and easily hydrolyzes to the more stable phases, OCP and HA. Lattice water molecules in DCPD structure are interlayered between the CaP chains arranged parallel to each other (Fig. 4). DCPD has been proposed as an intermediate in bone mineralization and in enamel dissolution [29]. Its conversion in anhydrous dicalcium phosphate (DCPA, CaHPO4) occurs at relatively low temperature [3]. DCPA, the mineral monetite, is  the anhydrous form of DCPD. DCPA crystallizes in the triclinic P1 space group. Its occurrence has never been reported in either normal or pathological calcifications. 2.7. Tetracalcium phosphate Tetracalcium phosphate (TTCP, Ca4(PO4)2) has the highest ratio Ca/P = 2. It crystallizes in the monoclinic P21 space group. Because

2.8. Calcium-deficient hydroxyapatite Calcium-deficient hydroxyapatite (CDHA, Ca10 x(HPO4)x(PO4)6 x (OH)2 x) has a Ca/P molar ratio less than the stoichiometric value of 1.67, and can be prepared either by precipitation or hydrolysis of ACP, DCPD, DCPA or a-TCP [8,15]. The structure contains vacant Ca ion sites and OH ion sites, whereas some of the PO34 ions are either protonated or substituted with other ions. In biological systems CDHA always occurs with ionic substitutions [10]. The XRD pattern is characteristic of a poorly crystalline apatite. The calculated powder XRD patterns of the most common CaPs are reported in Fig. 5. The numerous biomedical applications of CaPs include repair of bone and periodontal defects, scaffolds for regenerative medicine, coatings of metallic implants, bone cements, delivery systems for drugs, and antibacterial agents [8]. In addition, CaP nanoparticles have been proposed as second-generation vectors for gene therapy based on DNA transfection [30,31]. 3. Ionic substitutions in CaPs

Fig. 4. DCPD structure view along the a-axis. The phosphate ions are located between Ca ions arranged on pleated layers (highlighted by the black line connecting the Ca atoms). The water molecules fill layers between the Ca and phosphate ions.

Substitution of foreign ions in the crystal lattice of CaPs occurs frequently. In particular, the high stability and flexibility of the apatite structure accounts for the great variety of possible cationic and anionic substitutions. As a matter of fact, HA is just a member of a family of inorganic crystalline compounds, of general formula M10(XO4)6Y2. M is usually a bivalent cation, such as Ca2+, Sr2+, Ba2+, Cd2+, Pb2+, but monovalent and trivalent cations, such as Na+, K+ and Al3+ can be hosted as well; XO4 is usually PO34 ; VO34 or AsO34 , 4 but possible substitutions include also SiO4 ; CO23 and SO24 ; Y is a monovalent anion, OH , F , Cl , Br [1,3]. Cationic substitutions can occur for the whole range of composition, as happens for Sr2+, Cd2+ and Pb2+, or to just a limited extent as in the case of smaller ions,

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Fig. 5. X-ray diffraction patterns of the most common calcium phosphates (calculated using Cu Ka radiation k = 0.15418 nm).

which inhibit the crystallization of HA, such as Zn2+ and Mg2+ [32–34]. The biologically most relevant anionic substitutions involve F and Cl for OH , and CO23 for PO34 or OH group. Both F and Cl can replace OH over the whole composition range. The substitution of F is facilitated by its H-bonding properties, and it yields increased stability and crystallinity. A small amount of F substituted for OH drives the building of very long rows of iso-oriented OH . The resulting fairly well-ordered apatitic structure justifies the possible lower solubility of partially F-substituted HA compared to HA or fluoroapatite itself [35]. CO23 can substitute for PO34 (B substitution) or for OH (A substitution). A and B carbonated apatites can be distinguished on the basis of their different lattice constants, and by the different positions of the CO23 infrared absorption bands. In biological apatites, CO23 substitutes mainly for PO34 [36] in B-type apatite. Charge compensation by a Ca vacancy, together with a H atom which bonds to a neighboring PO34 , has been calculated as the most stable arrangement [37]. CO23 inhibits HA crystal growth, and its incorporation usually results in poorly crystalline structures with increased solubility [15,17]. The opposite effect of CO23 and F on crystallinity and solubility has been recently exploited to prepare apatites with tunable solubility. Structural investigation of the apatites at different ionic concentrations indicated that the partial substitution of CO23 for PO34 was balanced by the partial substitution of Na+ for Ca2+. In vitro dissolution experiments showed that increasing the CO23 and Na+ concentrations increased apatite solubility, whereas increasing the F concentration decreased it [38].

CO23 , as well as Mg2+ and F , were reported to suppress the growth of OCP along the c-axis direction under most conditions [39], and several bivalent and trivalent cations have been found to inhibit the growth of DCPD [40]. However, the hydrated phases OCP and DCPD are less sensitive than HA to the presence of impurities. Thus, the results of a seeded constant composition study demonstrated that Mg2+, which usually inhibits HA crystal growth, has a modest influence on the kinetic of OCP growth, and almost no effect on DCPD crystallization [41]. This difference in inhibitory effect has been suggested to be due to the presence of lattice water in OCP and DCPD, which decreases the adsorption of foreign ions [15]. However, the results of theoretical studies based on first-principles calculations, with OCP assumed to be a structural model of the transition region between HA and the solution, indicated that substitutional foreign ions, such as Sr2+, Mg2+ and Zn2+, can be more easily incorporated in OCP than in HA lattice [42,43]. Several ions have been reported to be able to replace, at least partially, Ca ions in the structures of a-TCP and b-TCP. In particular, ions with a smaller ionic radius than Ca2+, such as Mg2+, Mn2+ and Zn2+, usually stabilize the crystal lattice of b-TCP, whereas larger ions, such as Sr2+, can be accommodated more easily into a-TCP structure [44–48]. The presence of foreign ions can affect the hydrolysis of the more soluble CaPs into the more stable phases. Inhibition provides kinetic stabilization of initial transient phases, whereas hydrolysis promotion prevents detection of intermediate phases [15,40, 49,50]. Even the highly unstable ACP may persist for appreciable periods in the presence of Mg2+ ions, which kinetically stabilizes acidic CaP phases during in vitro precipitation reactions [51]. Not only the relative rate of crystallization and hydrolysis, but also the composition and structure of the CaP eventually formed, depend on the presence of foreign ions. Thus, the product of the hydrolysis can be a substituted phosphate, such as K+- and CO32 -substituted HA formed by the hydrolysis of OCP or of DCPD, and Sr2+-substituted HA obtained by hydrolysis of a-TCP [48,52–54]. 4. Recent developments in biological relevant ionic substitutions in CaPs According to the biomimetic principle that a biomaterial should resemble as closely as possible the host tissue, the optimum CaP for biomedical application should consist of apatite characterized by small dimensions, low crystallinity and non-stoichiometric composition. Thus, it is not surprising that HA is the most studied CaP, or that most research on ionic substitutions in CaPs involves HA. A biomimetic HA should also contain significant amount of CO23 ions. On this basis, several attempts have been performed in order to synthesize biomimetic nanocrystalline apatite containing carbonate as raw material for the manufacture of biomaterials [28,55]. In particular, Tadic et al. [28] proposed a method for preparing unlimited amounts of carbonated apatite with adjustable CO23 content, crystallinity and morphology. More recently, a synthetic route for the preparation of high amounts of nanocrystalline and nanometric carbonated apatite with a very high specific surface area has been reported [55]. Moreover, a significant part of the investigations on ionic substitutions in low-temperature CaPs have been performed on materials synthesized in air, which implies absorption of atmospheric CO2 on the crystallization products, or even in the presence of reagents that provide CO23 ions in the reaction medium, such as NaHCO3. It follows that most of the synthesized ion-substituted CaPs contains CO23 ions as well, and their chemical, structural and morphological properties are influenced by the co-presence of (at least) two foreign ions. The recent literature shows that several ions have attracted and attract remarkable interest related to their possible employment in the preparation of CaP-based biomaterials. Herein we will illus-

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trate the main results obtained for these ions, as well as those obtained in some application fields, which have explored the possibility of introducing ionic substitution as a tool to improve the properties of the developed materials. 4.1. Silicon Si is known to be essential as a trace element in biological processes. In particular, it has been reported to have a specific metabolic role connected to bone growth [56,57]. As a consequence, Si-substituted CaPs have attracted the interest of many scientists because Si incorporation is considered to be a promising way to improve the bioactivity of CaP-based biomaterials. The studies on material synthesized via an aqueous precipitation reaction are limited to Si-substituted HA [58] or to apatite/ACP mixtures [59], whereas most results have been derived from the products of sintering of apatites precipitated at low temperature [60–62]. The amount of Si which can be incorporated into HA seems to be limited to a maximum of 5 wt.% [60,63], and amounts around 1 wt.% have been suggested to be enough to elicit important bioactive improvements. Si plays a very important role at the microstructural level since it is responsible of the increase of the amorphous phase as well as of the reduction of the HA mean crystallite size (i.e. the mean size of coherently scattering domains) [59]. Si substitution influences the material surface by generating a more electronegative surface, by creating a finer microstructure, and then by increasing solubility [62]. The increased bioactivity of Si-substituted HA with respect to HA could be related to the increased number of defects, in particular those involving grain boundaries, that have been suggested to be the starting point of in vivo dissolution [64]. The simplest model for Si incorporation in CaPs is substitution 4 of SiO4 for PO34 groups, which is a non-isoelectronic substitution. To avoid a large cost in energy, local charge neutrality is mandatory and some other defect has to be associated with the PO34 substitution to compensate for the charge deficit. There are many possible mechanisms for charge compensation, such as oxygen or anionic vacancies, Ca and/or H excess—the predominant one depends on the thermodynamic conditions during the preparation of the materials. Unfortunately XRD is not an accurate enough technique to distinguish between phosphorus and Si since they are almost isoelectronic atoms, and the presence of H atoms cannot be usually detected. Neutron diffraction should give better insight into Si and H positions. However, wet preparation methods yield products with broad reflections, so that thermal treatment of the phosphate material is needed to obtain deeper structural details. Powder XRD pattern refinements (Rietveld method) of Si-containing HA heat treated at 1200 °C showed a contraction of the a-axis together with an increase in the c-axis and tetrahedral distortion [58]. However, the analysis was carried out on samples whose structure was quite far from that of the as-prepared samples. Solid-state nuclear magnetic resonance spectroscopy should be a powerful tool to 4 localize the SiO4 moieties and to investigate their environment also in low-crystalline materials. Recent results have been obtained on HA containing 4.6 wt.% Si. By using a specific editing sequence, the analysis unambiguously demonstrated that only a fraction of the Si atoms is incorporated into the HA lattice in the 4 form of SiO4 ion. A large amount of silicate units are located outside the HA structure and correspond to silica gel units [63]. The simultaneous incorporation of CO23 or Mg together with Si has been proposed in order to obtain materials very similar to biological-like apatites [59,65,66]. At low concentration, CO23 ions compensate for the deficit induced by Si, although they compete for the occupation of the PO34 site. Both CO23 and silicate reduce HA crystallinity, and the structure can host only a limited amount of the two ions before collapse [65].

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It has been reported that in vitro Si enhances and stimulates osteoblast-like cell activity [58] and in vivo induces a higher dissolution rate.[64]. Moreover, several studies suggest better biological activity of Si-substituted CaPs with respect to unsubstituted CaPs [62,67–69]. However, a recent analysis of the published results on Si-substituted CaPs correctly highlights the lack of experimental evidence that (i) Si ions are released from Si-substituted CaPs at therapeutic concentrations; and (ii) the improved biological performance of Si-substituted CaPs is indeed due to Si release, and not to Si-induced chemico-physical modifications of the CaP, such as superficial chemistry, topography and microstructure [70]. The same objections could be advanced to most of the papers claiming the promising biological activity of ion-modified CaPs, as it will appear in the following sections. 4.2. Strontium Sr has a great affinity for bone, where it is present in significant amounts, especially at regions of high metabolic turnover [71]. The beneficial effect of low doses of Sr in the treatment of osteoporosis has long been known [72], and its administration as Sr ranelate has recently been shown to reduce the incidence of fracture in osteoporotic patients [73,74]. Sr ranelate decreases bone resorption by inhibition of osteoclast-resorbing activity and osteoclastic differentiation, and it promotes bone formation by enhancing pre-osteoblastic cell replication and osteoblastic differentiation. The growing evidence of positive results obtained in clinical studies on long-term Sr treatment [75] has increased interest towards its incorporation in CaPs. Sr (ionic radius 0.12 nm) can replace Ca (ionic radius 0.099 nm) in the structure of HA over the whole range of composition, causing a linear expansion of the lattice constants. A detailed structural and morphological investigation carried out on HAs precipitated from aqueous solutions at different Sr concentrations demonstrated that relatively low Sr replacement of Ca induces a decrease in the coherent length of the perfect crystalline domains and disturbs the shape of the crystals, whereas the crystallinity, as well as the mean dimensions of the crystals, significantly increase at relatively high Sr content (Fig. 6) [76]. Moreover, the results of the structural refinements indicate that Sr exhibits a slight preference for the M(II) site—as expected on the basis of its greater ionic radius than that of Ca—for most of the composition range, but at very low content it displays a modest preference for the smaller M(I) site. This finding, which has been recently confirmed by the results of Rietveld refinements of synchrotron XRD patterns [19], has been ascribed to the optimization of the metal–oxygen interaction [76]. Extended absorption fine structure (EXAFS) analysis suggests increasing lattice disorder and loss of OH with increasing Sr content [19], in agreement with the decrease of crystallinity induced by Sr substitution lower than 50 at.% verified by several studies [76–78]. Destabilization of crystal structure by the larger Sr atom has been invoked to justify the significant increase in solubility with Sr content verified by solid titration of Sr-substituted HA [79]. The increased solubility is further confirmed by the greater release of Ca in Hanks’ balanced synthetic fluid from Sr-substituted HA than from HA granules [80], and it is enhanced in the presence of co-substituted CO23 [81]. Moreover, the extent of precipitation from simulated body fluid (SBF) and from a CaP supersaturated solution has been reported to decrease on increasing Sr content in solution [82,83]. The solubility data indicate that the formation of Sr-HA is not thermodynamically favoured, which might seem inconsistent with the ion effect on bone mineral. However, precipitation from supersaturated solution yields nanocrystalline HA with a degree of Sr substitution depending on its concentration in solution [82–84]. The formation of Sr-substituted HA in seeded

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affect osteoclast proliferation, which reduces with increasing Sr content. 4.3. Magnesium

Fig. 6. TEM images of Sr-substituted HA nanocrystals containing (a) 18 at.% and (b) 71 at.% of Sr.

Sr-containing SBF has been interpreted as suggesting that nucleation may be easier and that Sr incorporation is kinetically driven at high supersaturation [84]. Anyhow, the precipitation of Srsubstituted HA could have been expected on the basis of the Ostwald rules of stages [15], and it is in agreement with the significantly higher Sr content of new compared to old bone [71]. Two mechanisms have been described for Sr incorporation in bone: ionic substitution and ion exchange [71]. Biological apatites can exchange ions with the surrounding fluids in order to fulfil their primary biological functions. On this basis, ion exchange has been proposed as a method for surface activation of CaP-based biomaterials [85]. Langmuir-type isotherms determined for Sr2+ interaction with not-carbonated and carbonated nanocrystalline HA indicate a greater ion exchange on carbonated apatite, which has been ascribed to its greater proportion of surface hydrated layer as compared to the bulk. The hydrated layer, which mostly involves bivalent ions and water, with HPO24 ions in non-apatitic environments [86], like the ‘‘hydrated layer” in OCP, would promote the uptake of foreign ions [85]. Sr uptake by ion exchange, which was found to amount to 1.24 mmol g 1 for carbonated apatite and to 0.75 mmol g 1 for HA, is largely reversible in a Ca-rich solution, except for samples matured for extended periods of time [86]. The high reversibility of the process should imply a fast Sr release in biological fluids from Sr-activated, apatite-based biomaterials. In vitro tests on as-prepared materials have been performed only on Sr-substituted HAs [87]. The results of the investigation carried out using osteoblast-like MG63 cells and human osteoclasts indicate that Sr influences bone cell behaviour in a dose-dependent way. Sr concentrations in the range 3–7 at.% significantly stimulate osteoblast activity and differentiation, as shown by the increased alkaline phosphatase activity, and collagen type I and osteocalcin production. Moreover, even 1 at.%, Sr substitution is sufficient to

The amount of Mg associated with biological apatites is higher at the beginning of the calcification process, and decreases with increasing calcification [88,89]. This ion plays a key role in bone metabolism, since it influences osteoblast and osteoclast activity, and thereby bone growth [90]. It is well known that Mg has a marked inhibitory effect on HA nucleation and growth, whereas it selectively stabilizes more acidic precursor phases [91,15]. Thus, it is not surprising that Mg substitution for Ca in the structure of HA occurs only over a limited composition range (up to about 10 at.%) [92,93]. Rietveld structure refinements of HA synthesized in aqueous solutions with increasing Mg concentrations indicate that Mg preferentially occupies the M(II) site. This was explained by the shorter metal–oxygen distances and the less defined, more flexible geometry of anion coordination provided by this site with respect to the M(I) site [92]. In line with its smaller ionic radius (0.065 nm), Mg substitution for Ca causes a reduction in the lattice parameters of HA. Moreover, the degree of crystallinity of Mgsubstituted HA decreases with increasing Mg content [92,94]. The reduction of crystallinity of precipitated HA on increasing Mg content prevents an accurate evaluation of the lattice parameters by conventional powder XRD techniques. Moreover, the simultaneous presence of other ions—such as CO23 , which can be incorporated during synthesis at low temperature, or co-precipitated in order to obtain an apatitic phase more similar to biological carbonated apatite [94,95]—makes it quite hard to establish the effective substitution of Mg in the HA lattice. The same problem is encountered with other ions that reduce HA crystallinity, and in most cases of co-presence of multiple ions [66,96,65]. It follows that the term ‘‘substituted” should be used only when supported by structural evidence of ion incorporation into the crystal structure. In spite of the limited achievable substitution, the total amount of Mg associated with the apatitic phases precipitated from aqueous solution can amount up to 30 at.% [92,93] with the excess being in the amorphous phase and/or on the crystal surface. Whatever its localization, Mg-doped HA displays increased solubility with respect to stoichiometric HA [95], which may be related to reduced crystallinity and/or increased surface hydration. In agreement, Mg reversible uptake by ion exchange has been reported to affect the surface hydration state of the sample [97]. Near-infrared (NIR) and medium-infrared (MIR) spectroscopic data indicated that the samples enriched with Mg2+ retained more water at their surface than the Mg-free sample, both at the level of H2O coordinated to cations and adsorbed in the form of multilayers. Additionally, the H-bonding network in defective subsurface layers was also noticeably modified, indicating that the Mg2+/Ca2+ exchange involved was not limited to the surface [97]. Implantation of granules of 5.7 at.% Mg-doped HA in rabbit femoral bone defects was reported to enhance osteoconductivity and resorption with respect to commercial stoichiometric HA granulate, taken as control, in agreement with the greater solubility of Mg-doped HA [95]. 4.4. Zinc Zn is known to be an essential trace element for life. It functions as a cofactor for numerous enzymes, including those involved in DNA and RNA replication and protein synthesis [98]. Likewise, it is an important mineral in the normal growth and development of the skeletal system and its deficiency is associated with a decrease in bone density [99]. Zn inhibits osteoclast differentiation

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and promotes osteoblast activity [100], and thus promotes bone formation. Furthermore, a possible involvement of Zn has been suggested in the modulation of the morphology and crystallinity of biological apatite crystals [101]. Zn ion is one of the most effective of the simple metal ions in inhibiting HA crystal growth [32,102]. Nonetheless, it can be quantitatively incorporated into HA lattice by direct synthesis under mild conditions [103–105]. The substitution for Ca ions occurs up to about 20 at.%, whereas the apatite structure does not sustain higher Zn incorporation. The decrease in the lattice parameters observed up to 10 at.% substitution [105] is consistent with the smaller ionic radius of Zn2+ (0.075 nm) with respect to Ca2+ (0.099 nm). A correct evaluation of cell parameters at relatively high Zn content is hindered by the progressive broadening of the XRD peaks due to the reduction of crystal sizes and/or the increase of crystal strain with increasing Zn concentration [103]. Chemical analysis indicates that both Ca/P and (Ca + Zn)/P ratios decrease on increasing Zn content and, as a consequence, stoichiometry is no longer maintained. The loss in crystallinity is confirmed by scanning electron microscopy observations: apatite crystals display regular shapes, whereas the Zn-substituted apatite crystals are irregular and form agglomerates that grow in size with increasing Zn content. Zn exhibits an evident inhibitory role on the synthesis of HA through a reduction in crystallite size and a decrease in thermal stability [103,104]. In the presence of Zn ions, the constant composition growth of HA is markedly reduced and suggests a Langmuir-type adsorption at active growth sites on the crystals. Zn-containing HA particles have also been obtained by ion exchange in aqueous solutions at increasing Zn2+ concentrations. Zn uptake into HA crystals was found to reach about 2.7 at.% in 24 h [106]. Examination of the local coordination structure of Zn incorporated into HA by theoretical modelling gives evidence of the role of HA (0 0 0 1) face hydration in the Zn substitution process [107]. The energetics of Zn substitution and occupation on Ca(I) and Ca(II) sites showed that ion substitution is not favoured, but Zn occupation on Ca vacancy sites is highly probable (site M(II) is favoured) with spatial constraints being the most significant factor in the substitution mechanism. Density function theory calculations confirm that Zn prefers the M(II) site over the M(I) site, and favours tetrahedral coordination with significant local structural distortion [108]. Zn-containing apatite has also been obtained after hydrolysis of Zn-containing a-TCP powders [5]. The starting material was obtained as a single phase above 1300 °C. The higher the heating temperature, the more Zn is incorporated into a-TCP, with a maximum Zn content of 1.26 wt.%, obtained at 1450 °C. Soaking of Zn-containing a-TCP powders for 12 h in water resulted in a-TCP conversion with almost complete transfer of Zn from a-TCP to the apatite structure. 4.5. Manganese Mn influences regulation of bone remodelling, and its deficit causes reduction of organic matrix synthesis and retards endochondral osteogenesis, increasing the possibility of bone abnormalities such as decrease of bone thickness or length [109]. Mn supplement was found to be an effective inhibitor of loss of bone mass after ovariectomy [110]. Structural analysis of fluoroapatite synthesized in the presence of Mn2+ (ionic radius 0.090 nm) and heat treated at 900 °C indicates that Mn can substitute for Ca in the structure up to one atom per unit cell (10 at.% or 5.5 wt.%), and enters preferentially the M(I) positions, causing a slight rotation of the PO34 ion [111]. However, the presence of Mn2+ ion in solution has an inhibitory effect on HA crystallization and significantly reduces the amount of

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precipitate [109,112,83]. Furthermore, high concentration of Mn2+ at the surface of HA particles obtained by ion exchange was found to cause a rapid decrease in growth rate of HA in SBF solutions [109]. Nonetheless, HA containing small percentages of Mn can be easily prepared by precipitation from aqueous solutions, with a Mn concentration in the solid phase corresponding up to about 70% of that present in solution [112]. Apatites with low Mn content do not show any appreciable variation of the lattice parameters; however, Mn affects apatite thermal stability as evidenced by the partial transformation into b-TCP at 600 °C. Mn enters the structure of b-TCP more easily than that HA, as often occurs for small ions that inhibit HA crystallization, and indeed the lattice parameters of b-TCP obtained by thermal conversion of Mn-containing HA appear reduced with respect to those of the unsubstituted phase [113]. The morphology of Mn-doped apatites strongly depends on the ion concentrations. Transmission electron microscopy (TEM) images of samples containing both Mn and CO23 show that at relatively low Mn (0.73 wt.%) and high CO23 (3.6 wt.%) content large platelet and needle-like crystals were present, whereas at higher Mn (1.23 wt.%) and lower CO23 (2.1 wt.%) content only smaller needle-like crystals appeared [45]. It was suggested that the smaller and less perfect HA crystals with the higher Mn content should be more resorbable, and hence more biocompatible [45]. The coatings obtained on Ti substrates by pulsed laser deposition of these Mn-doped carbonated apatites were found to promote human osteoblast proliferation, activation of their metabolism and differentiation [114].

5. Applications in the biomedical field 5.1. Bone cements Calcium phosphate bone cements (CPCs) provide the biomimetic answer to the increasing demand of orthopaedic implant materials that can perfectly fit to a bone cavity, or that can be forged to the desired shape [115]. CPCs are constituted of one or several CaPs that, once mixed with a liquid phase, give a mouldable paste that stiffens during the setting reaction, which leads to the in situ formation of a solid CaP [116]. Different compositions yield two possible final products: apatite in the form of HA or CDHA, and DCPD [115]. As a consequence, cements are distinguished as either apatite CPCs and brushite CPCs. The involvement of biologically relevant ions in the composition of CPCs is aimed at improving both the setting properties and the biological performance of these materials. The effect of Sr2+ was tested both on brushite and on apatite CPCs. The brushite phase obtained from the setting reaction of a b-TCP/MCPA-based cement enriched with SrCl2 exhibits increased values of the lattice parameters in agreement with a partial substitution of Sr for Ca [117]. Sr displays a remarkable delaying action on the setting reaction, whereas it does not significantly influence the mechanical properties of brushite CPCs [117,118]. The introduction of Sr as Sr-substituted b-TCP leads to an increased presence of DCPA in the end products, which was ascribed to the lower pH values maintained for longer periods of time at higher Sr content. Release experiments under dynamic conditions for up to 15 days revealed the release of Sr2+ ions at dose ranges of 12– 30 ppm with zero-order release kinetics [118]. The addition of strontium hydrogen phosphate to the composition of TTCP-based cements has been reported to yield Sr-substituted HA, in agreement with the observed shift of the (0 0 2) reflection of HA on increasing Sr content [119]. The cements were submitted to in vitro and in vivo tests, and demonstrated good

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biocompatibility and increased degradation rates for both intramuscular and femoral implants. After 24 weeks in femur bone, the implants were not yet completely degradated; however, the highest degradation was found for the sample with the highest Sr content (10 at.%), which is in agreement with the effect of Sr on HA solubility [79]. Moreover, the results of EDAX analysis showed a gradient of Sr distribution in the interface between implant and new bone, which was suggested to support a possible role of Sr on new bone formation [120]. Sr-substituted HA was obtained as an end product of the setting reaction when Sr was added to the composition of ACP-DCPD-based cements as SrACP, as shown by the observed enlargement of the apatite lattice constants [121]. The end product consisted of two phases, HA and Sr-substituted HA when Sr was added as SrACP to ACP and DCPD, but only Sr-substituted HA, with a Sr content up to 10 at.%, when the composition contained only SrACP and DCPD [121]. The inhibitory effect of Sr on the hydrolysis reaction of a-TCP [48,54] can explain why Sr concentrations of just 5 at.% are sufficient to prevent setting of a-TCP-based cements [122]. However, at lower concentrations, the presence of Sr allows the setting times to be modulated, and displays a beneficial effect on bone cells, as shown by the increased alkaline phosphatase specific activity, collagen type I and osteocalcin production observed on cultures of osteoblast-like MG63 cells. Furthermore, the increased synthesis of osteoprotegerin indicates that the cements indirectly inhibit osteoclastogenesis and osteoclast function, in agreement with the results of in vitro tests on Sr-substituted HA nanocrystals [87]. The effect of Mg, Zn and F was explored on CPCs prepared with amorphous calcium carbonate phosphate (ACCP) doped with Mg, Zn or F ions [123]. The experimental CPCs consisted of a-TCP, doped ACCP and MCPM powders as matrix and biphasic calcium phosphate (BCP) granules. XRD analysis showed that the matrix converted to poorly crystalline apatite after setting for 24 h, while BCP remained apparently unchanged. The poor resolution of the XRD patterns did not allow verification of whether the ions had entered the crystal structure of HA. However, the doping ions influence the relevant properties of the cements: setting time and porosity were found to decrease in the order Mg > Zn > F, whereas the order is reversed for compressive strength [123]. The effect of Zn addition to CPCs on bone response was tested in vitro and in vivo using a-TCP with different levels of Zn substitution for Ca as cement powder, and sodium succinate solution as liquid phase [5,47]. The hydrolysis of a-TCP was significantly retarded by Zn incorporation greater than 0.26 wt.%, whereas no inhibition was observed for Zn concentrations lower than 0.11 wt.%. Furthermore, the results of in vivo tests carried out by implantation of the hardening bodies of a-ZnTCP cements at different Zn contents in the femora and tibia of white rabbits for 4 weeks indicated that Zn amount must not exceed 0.03 wt.% to avoid inflammation reaction [5]. CPCs possessing magnetic properties due to incorporation of iron oxides [124,125], and CPCs with high antimicrobial potency due to K and Na ions [126–128] have also been developed. 5.2. Biomimetic coatings Biomimetic methods for coating deposition involve mild conditions similar to those characteristic of biological environments, and allow deposition of CaP coatings on many different objects, such as sponges, cements, metal surfaces or fixation rods [129]. In particular, a thin layer of bone-like CaP deposited on metallic implants used for bone substitutions at high load-bearing sites offers a combination of the mechanical advantages of the metal with the excellent biocompatibility and bioactivity of the CaP. Since the first formulation of SBF solution proposed by Kokubo et al. [130], many efforts have been made to modify it and to improve the fast formation of adherent and homogeneous CaP layers. One of the main

advantages of biomimetic deposition of coatings is the possibility of co-precipitating ions, drugs, macromolecules and biological molecules together with the inorganic layer. Many studies have been addressed to investigate the influence of hydrogen carbonate concentration on the deposition from SBF. Fourier transform infrared spectroscopy, Raman spectroscopy and X-ray analyses of the layers precipitated from SBF at different HCO3 content indicate that it acts as a crystal growth inhibitor and influences the composition and structure of CaPs. It was suggested that as long as the HCO3 concentration is lower than 20 mmol l 1, only B-type carbonated HA precipitates, whereas at higher HCO3 concentrations, both A- and B-type form [131]. X-ray and selected-area diffraction TEM results indicated a preferred crystal growth orientation along the c-axis direction and perpendicular to the metal surface [132]. The deposition kinetics was found to be influenced by the HCO3 content also when the experiments were performed using SBF concentrated by a factor of 5 (SBF  5). It was found that HCO3 induces a reduction of apatite crystal size and so a better physical attachment on the metallic substrate [133]. In similar experimental conditions, it was verified that Mg2+ has an inhibitory effect even greater than that of HCO3 [134]. Nonetheless, Mg seems to play an important role at the coating/substrate interface. The relatively high Mg concentration at the interface shown by X-ray photoelectron spectroscopy was suggested to promote heterogeneous nucleation of tiny apatitic globules onto the substrate, with a consequent enhancement of physical adhesion at the early stage of the coating formation [134]. Mg was found to inhibit apatite crystallization also from a slightly supersaturated CaP solution. The reduction of the amount of precipitate induced by Mg is intermediate between that of Sr and that of Mn [83]. Mg incorporation in the apatitic phase up to 2 at.% provokes a shift of the (0 0 2) X-ray reflection, suggesting a partial substitution of Mg for Ca. An opposite shift of the (0 0 2) reflection was recorded for Sr-containing apatite deposited from CaP as well as from SBF, in agreement with a partial substitution of the larger Sr ion [82,83]. The ion that was demonstrated to be the strongest inhibitor of precipitation from the CaP calcifying solution is Mn2+ [83]. This ion remarkably hinders the deposition of HA and gives an amorphous phosphate relatively rich in Mn (up to about 30 at.%). In spite of the very small amount of precipitate and the thin layer that is produced, the Mn-containing coating displays a remarkably beneficial effect on bone cells [83]. The comparative in vitro investigation carried out using human osteoblast MG-63 on CaP coatings deposited on Ti in presence of Sr, Mg and Mn indicated that the Mg2+ and Sr2+ apatitic coatings promote proliferation and expression of collagen type I, with respect to bare Ti and to the thin layer of amorphous phosphate obtained in the presence of Mn2+. However, the relatively high content of Mn2+ in the phosphate has a remarkable beneficial effect on osteocalcin production, which is even greater than that observed for Sr2+. The significant effect of this ion on osteocalcin production was supported by the data obtained in parallel on osteoblast cultured in Dulbecco’s Modified Eagle’s Medium supplemented with increasing amounts of Mn [83]. Good in vitro results have also been obtained in the case of Zncontaining apatite layers deposited on titanium fixation rods [135]. The use of ZnCl2-containing supersaturated CaP solutions allowed the behaviour of the poorly crystalline apatite coatings with varying Zn concentration to be studied. During apatite formation, Zn is co-precipitated together with apatite so that an increase in Zn concentration in solution provokes an increase in the amounts of Zn and a decrease in Ca and phosphorus content on the fixation rod surface coating. Zn inhibits crystal growth, and the resulting low crystallinity of apatite makes it unclear whether or not substitution of Zn for Ca in the apatite crystal occurs. The Zn-containing layer is extremely thin (tenths of nanometers) and highly adherent to the

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metal surface. A significant increase in fibroblastic proliferation, osteoblastic proliferation and differentiation in vitro was observed as Zn precipitated on the Ti external fixation rods increased from 0 to 0.195 ± 0.020 lg cm 2 [135]. 5.3. Fluorescing markers Lanthanide-doped CaP nanoparticles show very interesting luminescent properties which make them useful for biological fluorescent labelling. Their fluorescence is characterized by narrow emission band widths determined by the lanthanide ions, high photochemical stability and long fluorescence lifetimes (up to several milliseconds). The orbital energy structure in lanthanide ions also offers the possibility of large shifts between the excitation and emission bands. Europium ions can co-precipitate with Ca ions during synthesis with a maximum Eu/(Eu + Ca) ratio of about 2 at.% [31,136–138]. Europium-doped apatite with a mean particle size of 620 nm, resulting from a broad size distribution (between 475 and 950 nm) has been obtained in alcoholic solution at physiological temperature [136,137]. The most efficient emission of europiumdoped nanoparticles can be observed at 614 nm; as a consequence the luminescence in the visible region is typically red. Changing the nature of the luminescent centre of lanthanide-doped CaP nanoparticles will result in sharp line emission spectra with peak maxima varying over a large spectral range (from red to blue). Colloidal dispersions stable at neutral pH and emitting red or green colours under excitation at 253.7 or 365 nm were obtained for Eu3+ and Tb3+, respectively [138]. Lanthanide-doped CaP nanoparticles have also been prepared in aqueous solutions. Variation of the temperature of synthesis allows the crystallinity of the nanoparticles to be controlled, from amorphous at 37 °C to nanocrystalline HA at 80 °C [31]. Stable colloids of Eu-doped and Tb-doped HA nanoparticles were obtained by functionalization with DNA. In this context lanthanide-doped CaP nanoparticles have shown great potential for cell transfection

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applications since observations on living cells are possible as emission is in the visible region of the spectrum. The fluorescence efficiency depends strongly on the internal crystallinity: amorphous nanoparticles show much weaker fluorescence than the crystalline ones. However, the nanocrystalline samples synthesized at 80 °C conjugate a sufficiently intense fluorescence with biologically suitable particle size (Fig. 7) [31]. The studies on lanthanide-doped CaP nanoparticles do not present direct evidence that the lanthanide ions are substituted in the structure of apatite, due to the low content of doping ion and/or low crystallinity of the apatitic phase. However, ionic substitution has been hypothesized on the basis of the observed luminescence. In fact, the absence of luminescence was reported when the doping lanthanide ions were only adsorbed on the surface and not incorporated in the apatitic structure, whereas the samples became luminescent after heating at 800 °C, as the rare earth ions diffused into the apatitic lattice [139]. 5.4. Biomagnetic nanoparticles Magnetic materials are relevant in biology and biomedicine because of their responses to applied magnetic fields. In particular, CaP nanoparticles with magnetic properties may find applications as heating mediators in hyperthermia therapy for cancer, a technique that has recently attracted increasing attention [140]. Magnetic HA and DCPD nanoparticles have been recently prepared by co-precipitation in aqueous solution with the addition of different concentrations of Fe2+ [141,142]. The magnetic properties of these nanoparticles are influenced by several factors, such as microstructure, anisotropy and orientation. For this reason, an excess of iron should be avoided since it does affect the arrangement of atoms in the CaP lattice and damages the integrity of the crystal structure. In fact, it is quite hard to establish if iron is really substituted for Ca in the magnetic crystal lattices. The presence of Fe2+—which has a distinctly smaller radius (0.0835 nm) than Ca2+—should clearly affect the lattice parameters. However, published results are some-

Fig. 7. Optical micrographs (transmission mode) (left), confocal laser scanning micrographs (centre) and overlap of both micrographs (right) of the Eu-doped HA (excited at 458 nm) (a) and Tb-doped HA (excited at 364 nm) (b) nanoparticles. The colours in the overlap images (right) were slightly dimmed in order to increase the contrast [31]. Reproduced by permission of The Royal Society of Chemistry.

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what confused about this matter. It is clear that the presence of iron in association with CaPs confers peculiar magnetic properties on the nanocrystals, which are useful in practical application. In vitro tests of hyperthermia therapy showed encouraging results, suggesting that therapy conducted by magnetic DCPD could kill cancer cells without harming normal cells [142]. Magnetic HA nanoparticles have also been tested in vivo by injecting around the tumour magnetic HA nanoparticles or pure HA nanoparticles mixed with phosphate buffer solution (PBS). Tumours injected with iron-containing HA and treated inside the magnetic field to achieve hyperthermia showed dramatic reduction of volume in the 2 week observation period [143].

6. Concluding remarks The analysis of the data reviewed in this paper provides evidence of the growing interest towards ionic substitution as a powerful tool to improve the biological performance of CaPs, either through modifications of the structural, morphological and chemical characteristics of the CaP or even through exploitation of the therapeutic properties of the substituting ion. Most of these researches have been focused on apatite. On the other hand, it is evident that a number of other CaPs are finding widespread employment in biomedical applications. It follows that the extension of the structural investigation on the role of ionic substitutions to further phases should yield remarkable useful information for improving and modulating the properties of CaP-based materials. In most cases the available data do not provide sufficient information to establish if the biological response is due to the direct effect of the ion or to the CaP modifications, such as changes in superficial chemistry and morphology induced by the ion. The following critical points emerge from our overview of the published results: – The term ‘‘ion-substituted” is quite often used without any experimental proof of the regular incorporation of the ion inside the crystal lattice of the CaP. The difficulty generally arises from the low degree of crystallinity of the inorganic phase, which may be due to several factors, including (i) the conditions of preparation—biomimetic syntheses are usually performed at low temperature; (ii) inhibition of crystal growth due to the presence of the foreign ion, which destabilizes the crystal structure; (iii) co-presence of further ions—incorporation of CO23 ion cannot be avoided at low temperature, and its co-presence is often desired in order to obtain a product more similar to biological carbonated apatites. It follows that many ‘‘substituted CaPs” could indeed be ‘‘doped-CaPs”, where the foreign ion is just adsorbed on the surface of the crystals. – Crystallinity and crystal dimensions significantly affect solubility and, as a consequence, ion release. Thus, a decrease in structural order and/or increase in superficial area due to the presence of the foreign ion might be responsible of the observed increase in solubility, without implying an active role of the ion. – Ion release proceeds much more quickly when the ion is just adsorbed on the crystal surface than when it is inside the crystal structure. The rate and amount of ion release are very important properties for biomedical applications; nonetheless only a few papers have investigated ion release in synthetic fluids [64,80,81,85,86,118], and even fewer in in vivo experiments [120]. – Only a minor part of the developed ion-substituted and/or iondoped CaPs have been submitted to in vitro and/or to in vivo tests, and the improvement of the biological response due to the ion is thus quite often just an hypothesis. Moreover, a CaP with a different superficial chemistry, topography and micro-

structure from that of the ion-modified CaP is usually utilized as a control material for biological tests, thereby preventing a correct analysis of the results. Only in a few cases the modified in vitro or in vivo biological response can be ascribed to the presence of the foreign ion [83,87,120]. The above considerations highlight the general need of more detailed analyses of ion-substituted CaPs in terms of (i) modifications of the structural, morphological and chemical characteristics of the CaP induced by the ion; (ii) in vitro and in vivo ion release; (iii) biological response based on in vitro and in vivo tests performed using a suitable control material. Nonetheless, this is an exciting challenge in view of the promising future of ion substitution for the development of CaP materials with improved bioactivity, which can be also used as delivery systems of potentially therapeutic ions, suggested by the recent achievements reviewed in this paper. Appendix A. Figures with essential colour discrimination Certain figures in this article, particularly Figures 1, 2, 4, 5 and 7 are difficult to interpret in black and white. The full colour images can be found in the on-line version, at doi:10.1016/j.actbio.2009. 12.041. References [1] Kay MI, Young RA, Posner AS. Crystal structure of hydroxyapatite. Nature 1964;204:1050–2. [2] Dorozhkin SV. Calcium orthophosphates in nature, biology and medicine. Materials 2009;2:399–498. [3] Elliott JC. Structure and chemistry of the apatites and other calcium orthophosphates. Amsterdam: Elsevier; 1994. [4] McConnell D. Apatite. Its crystal chemistry, mineralogy, utilization, and geologic and biologic occurrences. Wien: Springer-Verlag; 1973. [5] Li X, Sogo Y, Ito A, Mutsuzaki H, Ochiai N, Kobayashi T, et al. The optimum zinc content in set calcium phosphate cement for promoting bone formation in vivo. Mater Sci Eng C 2009;29:969–75. [6] Garcia F, Ortega A, Domingo JL, Corbella J. Accumulation of metals in autopsy tissues of subjects living in Tarragona County, Spain. J Environ Sci Health 2001;36:1767–86. [7] Carlisle EM. Silicon: a possible factor in bone calcification. Science 1970;167:279–80. [8] LeGeros RZ. Calcium phosphate-based osteoinductive materials. Chem Rev 2008;108:4742–53. [9] Weiner S, Wagner HD. The material bone: structure–mechanical function relations. Annu Rev Mater Sci 1998;28:271–98. [10] Dorozhkin SV, Epple M. Biological and medical significance of calcium phosphates. Angew Chem Int Ed 2002;41:3130–46. [11] Onuma K, Ito A. Cluster growth model for hydroxyapatite. Chem Mater 1998;10:3346–51. [12] LeGeros RZ. Properties of osteoconductive biomaterials: calcium phosphates. Clin Orthop 2002;395:81–98. [13] Xin R, Leng Y, Wang N. In situ TEM examinations of octacalcium phosphate to hydroxyapatite transformation. J Cryst Growth 2006;289:339–44. [14] Suzuki O, Kamakura S, Katagiri T, Nakamura M, Zhao B, Honda Y, et al. Bone formation enhanced by implanted octacalcium phosphate involving conversion into Ca-deficient hydroxyapatite. Biomaterials 2006;27:2671–81. [15] Wang L, Nancollas GH. Calcium orthophosphates: crystallization and dissolution. Chem Rev 2008;108:4628–69. [16] McDowell H, Gregory TM, Brown WE. Solubility of Ca5(PO4)3OH in the system Ca(OH)2–H3PO4–H2O at 5, 15, 25, and 37 °C. J Res Natl Bur Stand A Phys Sci 1977;81A:273–81. [17] Featherstone JDB, Mayer I, Driessens FCM, Verbeeck RMH, Heijligers M. Synthetic apatites containing Na, Mg, and CO3 and their comparison with tooth enamel mineral. Calcif Tissue Int 1983;35:169–71. [18] Bigi A, Gazzano M, Ripamonti A, Foresti E, Roveri N. Thermal stability of cadmium–calcium hydroxyapatite solid solutions. J Chem Soc Dalton Trans 1986:241–4. [19] Terra J, Rodrigues Dourado E, Eon JG, Ellis DE, Gonzalez G, Malta Rossi A. The structure of strontium-doped hydroxyapatite: an experimental and theoretical study. Phys Chem Chem Phys 2009;11:568–77. [20] Brown WE, Smith JP, Lehr JR, Frazier AW. Octacalcium phosphate and hydroxyapatite. Nature 1962;196:1050–5. [21] Bodier-Houllè P, Steuer P, Voegel JC, Cuisinier FJC. First experimental evidence for human dentine crystal formation involving conversion of octacalcium phosphate to hydroxyapatite. Acta Cryst 1998;D54:1377–81.

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