**10.10. Collagen apatite composites**

Tissue engineering techniques have been developed to recover or enhance lost tissue func‐ tion and structure.12 Collagen-apatite13 (or collagen/apatite, **Col-AP**) composite resembling the composition of natural bone10 has been studied extensively and considered as a promising bone tissue engineering material, which can be used to replace or regenerate damaged tissue, resulting from an accident, trauma or cancer. Such synthetic or hybrid biomaterials must have high porosity with interconnected pores to allow the vascularization as well as the nutrients and gases diffusion. Moreover, they should be biodegradable to act as temporary cellular support [114],[126],[127],[128].

The defining feature of collagen is a structural motif in which three parallel polypeptide strands in a left-handed, polyproline II-type (**PPII**) helical conformation coil around each other with a one-residue stagger to form a right-handed triple helix. Tight packing of **PPII** helices within the triple helix mandates that every third residue be **GLY**, resulting in a repeating **XAAYAAGLY** sequence, where **XAA** and **YAA** can be any amino acid. This repeating occurs in all types of collagen, although it is disrupted at certain locations within the triple-helix domain of nonfibrillar collagens. The amino acids in the **XAA** and **YAA** positions of collagen are often (2S)-

<sup>12</sup> One of the first polymers used for bone tissue engineering was based on a hydrolytically copolymer of polylactic-coglycolic acid (PLGA) but its use for large bone defect regeneration was controversial as inflammatory events were observe. The utilization of chitosan and alginate was also investigated [128].

<sup>13</sup> Names such as collagen-hydroxylapatite or collagen-hydroxyapatite composite were also often applied in published literature.

proline (**PRO**, 28%) and (2S,4R)-4-hydroxyproline (**HYP**, 38%), respectively. The **PROHYPGLY** is the most common triplet (10.5%) in collagen (**Fig. 18**) [129].

The age variations of the crystal lattice parameters of human enamel apatites are related to complicated processes of de- and remineralization, which result in the increase or reduction

( ) [ ] ( ) [ ]

n m/ 2 n A m ,H O -

+- - -

values of *a* and *c*-parameter of enamel apatites change considerably without any depend‐ ence of particular age that may be explained by essential fluctuations of the content of Ca in human organism. After 50 years of age, significant direct correlation between the age and the

The surface of apatite nanocrystals is possibly doped with foreign elements or functional‐ ized with organic molecules [117],[123],[124]. The course of facile synthesis of B-type carbo‐

Tissue engineering techniques have been developed to recover or enhance lost tissue func‐

composition of natural bone10 has been studied extensively and considered as a promising bone tissue engineering material, which can be used to replace or regenerate damaged tissue, resulting from an accident, trauma or cancer. Such synthetic or hybrid biomaterials must have high porosity with interconnected pores to allow the vascularization as well as the nutrients and gases diffusion. Moreover, they should be biodegradable to act as temporary cellular

The defining feature of collagen is a structural motif in which three parallel polypeptide strands in a left-handed, polyproline II-type (**PPII**) helical conformation coil around each other with a one-residue stagger to form a right-handed triple helix. Tight packing of **PPII** helices within the triple helix mandates that every third residue be **GLY**, resulting in a repeating **XAAYAAGLY** sequence, where **XAA** and **YAA** can be any amino acid. This repeating occurs in all types of collagen, although it is disrupted at certain locations within the triple-helix domain of nonfibrillar collagens. The amino acids in the **XAA** and **YAA** positions of collagen are often (2S)-

<sup>12</sup> One of the first polymers used for bone tissue engineering was based on a hydrolytically copolymer of polylactic-coglycolic acid (PLGA) but its use for large bone defect regeneration was controversial as inflammatory events were observe.

<sup>13</sup> Names such as collagen-hydroxylapatite or collagen-hydroxyapatite composite were also often applied in published

nated nanoapatite with tailored microstructure is described by GUALTIERI et al [125].

+ ++ (21)

2−, and []Ca, []OH− are the vacancies. Until the age of 50 years, the

(or collagen/apatite, **Col-AP**) composite resembling the

2 3 4 2 Ca OH 2

+ +®

n m / 2 Ca n PO m OH

2−, H2O and HPO4

2− contents

of vacancies in Ca positions and in the respective changes of CO3

486 Apatites and their Synthetic Analogues - Synthesis, Structure, Properties and Applications

( )

2− or HPO4

**10.10. Collagen apatite composites**

Collagen-apatite13

The utilization of chitosan and alginate was also investigated [128].

in the unit cell:

where A2− = CO3

tion and structure.12

literature.

support [114],[126],[127],[128].

*a*-parameter appears [122].

**Fig. 18.** The triple helix of collagen formed from (PROHYPGLY)4-(PROHYPALA)-(PROHYPGLY)5 (a), the view down the axis of a (PROPROGLY)10 triple helix (b) and the segment of triple helix (c) with hydrogen bonds (—) [129].

The categories of collagen14 include the classical fibrillar and network-forming collagens, the FACITs (fibril-associated collagens with interrupted triple helices), MACITs (membraneassociated collagens with interrupted triple helices) and MULTIPLEXINs (multiple triple-helix domains and interruptions). The collagen of type I is the most abundant type used in tissue engineering. Natural polymer collagen that represents the matrix material of bone, teeth and connective tissue can be extracted from animal or human sources (skin, bones, tendons, ligaments and cornea). The treatment includes the separation and isolation (in soluble or insoluble form), decalcification, purification (purification is required to eliminate the antigen‐ ic component of protein), sterilization and chemical modification process to achieve polya‐ nionic or purified protein. Type I polyanionic collagen was found to improve the cell adhesion [129],[130].

Collagen, the most abundant protein in extracellular matrix, is chemotactic to fibroblasts. It shows high affinity to cells and good resorbability in vivo. Nevertheless, its poor mechanical properties have restricted its usage in load-bearing applications. Carbonated apatite and collagen interact to form a composite material, the mechanical, physicochemical and biologi‐ cal properties of which differ considerably from those of either constituent considered separately. Collagen and non-collagenous proteins (**NCPs**) are thought to control the crystal deposition, size, crystallization and multiplication/maturation. On the other hand, the crystal deposition in intrafibrillary spaces is likely to modify the three-dimensional conformation of collagen [114],[131],[132],[133]. The solubility of apatite-collagen composites is significantly reduced by UV radiation [134],[135].

<sup>14</sup> There are 27 types of collagen described in literature composed of at least 46 distinct polypeptide (CRP, collagen-related peptide) chains [129], but the types I – V are the most common. More than 90% of collagen in human body is the fibrillar type I.

Calcium phosphates are available commercially, as hydroxylapatites are extracted from bones or they can be produced wet by direct precipitation from pH-adjusted solutions of calcium and phosphate salts [130]. The crystallographic c-axes of the plate-shaped apatite crystals are well aligned with long axes of collagen fibrils (**Fig. 19**), and this preferred orientation be‐ tween the mineral and the organic framework is assumed to be the general feature of the calcium phosphate biomineralization process. Several attempts were made to mimic this lowest level of hierarchical organization of bone by using proteins as a site for the heteroge‐ neous nucleation and subsequent growth of stoichiometric hydroxylapatite crystals. In special approaches, the biomimetic apatite coatings on surfaces were prepared by soaking the materials in simulated body fluid (SBF) solutions, which contained ions in the concentra‐ tions similar to those in inorganic part of human blood plasma. It is generally accepted that the in vitro apatite growth during the exposure to SBF is an indicator for the in vivo bioactiv‐ ity of materials surface [136],[137].

**Fig. 19.** The crystal structure of hydroxylapatite is oriented according to the c-axis with the extended direction of colla‐ gen fibrils [112].

Biomimetic materials are able to mimic the morphological and physicochemical features of biological apatite compounds, i.e. they are synthetic analogues of inorganic part of hard tissues [112],[117]. The biomimetic deposition is considered as an ideal method to produce calcium phosphate ceramics such as apatite coatings on titanium and its alloys for medical applications. It has also been proved that the chemical pretreatment in alkali solution can improve the bonding between the titanium substrate15 and calcium phosphate coatings fabricated by subsequent biomimetic deposition in simulated body fluid (SBF) [138],[139], [140],[141]. The biomimetic growth of apatite was also described on hydrogen-implanted silicon [142], polyvinyl alcohol (PVA) [143], TiO2 nanotubes [144], alumina [145], zirconia ceramics (Y-TZP) [146], forsterite [147], akermanite [148], magnetite [149], glasses [150] and

<sup>15</sup> Titanium and its alloys are widely used as orthopedic and dental implant materials because of their high mechanical strength, low modulus and good corrosion resistance [139].

bioglasses [151],[152], cements [153], geopolymers [154], carbon nanotubes [155] and micro‐ spheres [156].

Calcium phosphates are available commercially, as hydroxylapatites are extracted from bones or they can be produced wet by direct precipitation from pH-adjusted solutions of calcium and phosphate salts [130]. The crystallographic c-axes of the plate-shaped apatite crystals are well aligned with long axes of collagen fibrils (**Fig. 19**), and this preferred orientation be‐ tween the mineral and the organic framework is assumed to be the general feature of the calcium phosphate biomineralization process. Several attempts were made to mimic this lowest level of hierarchical organization of bone by using proteins as a site for the heteroge‐ neous nucleation and subsequent growth of stoichiometric hydroxylapatite crystals. In special approaches, the biomimetic apatite coatings on surfaces were prepared by soaking the materials in simulated body fluid (SBF) solutions, which contained ions in the concentra‐ tions similar to those in inorganic part of human blood plasma. It is generally accepted that the in vitro apatite growth during the exposure to SBF is an indicator for the in vivo bioactiv‐

> Hydroxyapatite (HAp) c

> > a

and calcium phosphate coatings

ity of materials surface [136],[137].

gen fibrils [112].

Collagen fibers

488 Apatites and their Synthetic Analogues - Synthesis, Structure, Properties and Applications

Extended direction

improve the bonding between the titanium substrate15

strength, low modulus and good corrosion resistance [139].

a

Ca

**Fig. 19.** The crystal structure of hydroxylapatite is oriented according to the c-axis with the extended direction of colla‐

Biomimetic materials are able to mimic the morphological and physicochemical features of biological apatite compounds, i.e. they are synthetic analogues of inorganic part of hard tissues [112],[117]. The biomimetic deposition is considered as an ideal method to produce calcium phosphate ceramics such as apatite coatings on titanium and its alloys for medical applications. It has also been proved that the chemical pretreatment in alkali solution can

fabricated by subsequent biomimetic deposition in simulated body fluid (SBF) [138],[139], [140],[141]. The biomimetic growth of apatite was also described on hydrogen-implanted silicon [142], polyvinyl alcohol (PVA) [143], TiO2 nanotubes [144], alumina [145], zirconia ceramics (Y-TZP) [146], forsterite [147], akermanite [148], magnetite [149], glasses [150] and

<sup>15</sup> Titanium and its alloys are widely used as orthopedic and dental implant materials because of their high mechanical

P

Ca10(PO4)6(OH)2 P O OH-

The ability to form apatite [150],[153] from supersaturated solution has been widely used to imply the bioactivity of an implant in vivo. However, the method itself may provide at best incomplete information, primarily because it is determined only by the solution supersatura‐ tion, irrespective of biological processes. The bone regeneration is triggered mainly by the vitality of osteoblasts and regulated by the expression of growth factors such as estrogen, parathyroid hormone and bone morphogenetic proteins, while ions or other species released from an implant may affect the expression of such growth factors and so the bone resorption or formation. The misinterpretation of the outcome of such tests must result in the misunder‐ standing of the true effects and behavior of materials intended for use in embedded biologi‐ cal applications. Moreover, SBF may not be able to mimic properly the physiological conditions because it is based on analytical concentrations and not on the activity of key components [157].

The methods of the preparation of collagen-hydroxylapatite composites include the produc‐ tion of composite gels, films, collagen-coated ceramics, ceramic-coated collagen matrices and composite scaffolds for bone substitutes and hard tissue repair via the following techniques [130],[158],[159]:

**• In vitro collagen mineralization**: the method is based on direct mineralization of a collagen substrate (film) through which calcium and phosphate ions diffuse into the fibrils or as phosphate-containing collagen solution16 (in situ precipitation [159]). **Fig. 20** introduces the experimental set-up for direct mineralization of a collagen sheet. The growth of HAP crystals with c-axis oriented along the collagen fibrils requires the pH in the range from 8 to 9 and the temperature of 40°C. These conditions promote the accumulation of calcium ions on the carboxyl groups of collagen molecules, leading to the nucleation of hydroxylapatite.

**Fig. 20.** Simplified scheme of apparatus for the preparation of collagen-hydroxylapatite composite by in vitro collagen mineralization method: modified scheme according to WAHL AND CZERNUSZKA [130] for in situ colla‐ gen/HAP precipitation (*a)*; modified scheme according to WANG and LIU [159]: pH meter (1), pH electrode (2), stirrer (3), peristaltic pump (4), hot-plate magnetic stirrer (5a) or stirrer (5b), thermocouple (6), hot plate (7) and peristaltic pump driven by pH controller (8) (b).

<sup>16</sup> These techniques usually continue with freeze-drying.


**Fig. 21.** The cycle of enzymatic mineralization of collagen sheets [130].


Under the condition of critical point,18 the density of liquid and gas phase converge and become identical (supercritical fluid) as well as the surface tension is negligible.

<sup>17</sup> Aqueous solution of sodium phosphate and chloride, where the ion concentrations and osmolarity (osmotic concen‐ tration) correspond to human body. In some cases, potassium phosphate and chloride were used.

<sup>18</sup> Critical point is defined by the value of three parameters including critical temperature, critical pressure and critical volume. Critical temperature is the highest temperature at which pure matter can exist as liquid.

**•** Col-HAP was cast into the mould and frozen, ice crystals replaced with ethanol, ethanolliquid CO2 exchanged and critical point dried to finally arrive with an exact porous replica of the original bone. Solid freeform fabrication techniques have recently been developed with artificial polymers and ceramic materials. These have the ability to change the pore interconnectivity, pore size and pore shape but have the disadvantage of not having the affinity of collagen to cell attachment. Another major advantage of Col-HAP scaffolds produced through the SFF method is the ability to control variables such as the control of external and internal structure, porosity and cross-linking [130].

**• Thermally triggered assembly of HA/collagen gels**: the solution of calcium and phos‐ phate ions is encapsulated within the liposomes and next inserted into the acidic suspen‐ sion of collagen. After the injection into a skeletal defect, the increasing temperature due to body heat initiates the gelation process, which leads to fibrous network. The mineraliza‐

**• Vacuum infiltration of collagen into a ceramic matrix**: ceramic scaffold is prepared by heating aqueous hydroxylapatite slurry containing poly(butyl methacrylate) (PBMA) spheres to high temperatures. The pyrolysis of PBMA particles leads to porous HAP green body. Pores in this matrix are then filled with collagen suspension under vacuum. The final

**• Enzymatic mineralization of collagen sheets**: is a method based on the cycle (**Fig. 21**) where collagen-containing alkaline (basic) phosphatase (ALP) is treated by aqueous solution of

phosphate to crystallize and the mineralization occurs on coated area. The sample is then coated again with collagen suspension, air-dried and cross-linked with UV irradiation. Repeating this cycle results in multilayered composite sheets of calcium/phosphate and

> Alkaline phosphatase in collagen material

**• Water-in-oil emulsion system**: purified collagen suspension mixed with HAP powders at the temperature of 4°C is next dispersed in olive oil and stirred at 37°C. The collagen aggregates and reconstitutes into the aqueous droplets. The addition of phosphate-buf‐

The main disadvantages of this method are the problem with complete removing of the oil

**• Freeze-drying and critical point drying (CPD) scaffolds**: the ice crystals with collagen fibers at the interstices can be formed by freezing the suspension of collagen and HAP in water under controlled conditions. In the case when the freeze-drying is applied (temperature and

<sup>17</sup> Aqueous solution of sodium phosphate and chloride, where the ion concentrations and osmolarity (osmotic concen‐

18 Critical point is defined by the value of three parameters including critical temperature, critical pressure and critical

(PBS) leads to the gelation of Col-HAP microspheres (gel beads) of bone filler.

3− ions for calcium

Collagen/ calcium phosphate composite

the density of liquid and gas phase converge and

Ca2+

tion occurs after reaching the liposome's transition temperature of 37°C.

490 Apatites and their Synthetic Analogues - Synthesis, Structure, Properties and Applications

composite was then freeze-dried to produce the microsponges within.

calcium ions and phosphate ester. The enzyme provides a reservoir for PO4

Alkaline

**Fig. 21.** The cycle of enzymatic mineralization of collagen sheets [130].

content from the composite and too low viscosity of the mixture.

pressure corresponding to CPD), ice crystals sublimate to water vapor.

tration) correspond to human body. In some cases, potassium phosphate and chloride were used.

volume. Critical temperature is the highest temperature at which pure matter can exist as liquid.

become identical (supercritical fluid) as well as the surface tension is negligible.

Glycerol 1 phosphate

Under the condition of critical point,18

collagen.

fered saline17

The collagen-hydroxyapatite/pectin (Col-HA/pectin) composite was prepared in situ by the introduction of pectin, a kind of plant polysaccharide, into the collagen-hydroxyapatite composite. The structure of composite consisted of hydroxylapatite of low crystallinity with particles uniformly dispersed in organic materials. There is strong bonding interaction between HAP, collagen and pectin. The mechanical properties, water absorption, enzyme degradation and cytotoxicity indicate a potential use in bone replacement for the new composite [160].

Most mammalian biofluids are supersaturated with respect to the bone and tooth mineral hydroxylapatite. Nevertheless, the biofluids, which are in contact with soft tissues, especial‐ ly those like milk that must be stored for any length of time, should be highly stable with respect to calcium phosphate precipitation. In contrast, saliva and the extracellular matrix of hard tissues, especially near to the sites of mineralization or remineralization, are required not only to maintain the mineral phase with which they are in contact but also to deposit calci‐ um phosphate in a highly controlled manner [161].

Collagen composites were also used as a scaffold for the repair of soft tissues. These materi‐ als provide analogous environment to extracellular matrix (ECM) and induced rate of synthesis or growth of new tissues. Several natural polymers as collagen, chitosan,19 gelatin20 and keratin21 possess the ability to induce the proliferation of cells and hence find their use as biomaterials for a wide range of biomedical applications. Among all biopolymers, collagen is a widely accepted material for the tissue engineering applications in the view of its low antigenicity, excellent biocompatibility and biodegradability [162].

The preparation of collagen-based biocomposite constructed from micro-crimped long collagen fiber bundles extracted from a soft coral embedded in alginate hydrogel matrix was described by SHARABI et al [163]. This biocomposite demonstrated the hyperelastic behavior similar to human native tissues.

<sup>19</sup> Chitosan is (1–4)-linked 2-amino-2-deoxy-b-glucan, a byproduct of N-deacetylation of chitin. It is a major constituent of crab and shrimp shells and of cuticles of insects.

<sup>20</sup> Gelatin is a denatured form of collagen. Gelatin has low antigenicity and promotes the cell adhesion, differentiation and proliferation. Gelatin also possesses high cytocompatibility, which makes it a potential candidate as a biomaterial for various tissue engineering applications [162].

<sup>21</sup> Keratin is a family of fibrous proteins, which is found abundantly in nature. It is the main constituent of hair, wool, nails, horns and hooves of mammals, birds and reptiles [162].
