Composite Films of Gelatin and Hydroxyapatite/Bioactive Glass for Tissue-Engineering Applications

Journal of Biomaterials Science 21 (2010) 1207–1226brill.nl/jbs

Composite Films of Gelatin and Hydroxyapatite/BioactiveGlass for Tissue-Engineering Applications

Piergiorgio Gentile a,∗, Valeria Chiono a, Francesca Boccafoschi a, Francesco Baino b,

Chiara Vitale-Brovarone b, Enrica Vernè b, Niccoletta Barbani c

and Gianluca Ciardelli a

a Department of Mechanics, Politecnico di Torino, Corso Duca degli Abruzzi 24, 10129 Turin, Italyb Department of Materials Science and Chemical Engineering, Politecnico di Torino,

Corso Duca degli Abruzzi 24, 10129 Turin, Italyc Department of Chemical Engineering, Industrial Chemistry and Materials Science,

University of Pisa, Via Diotisalvi 2, 56126 Pisa, Italy

Received 22 April 2009; accepted 9 July 2009

AbstractCross-linked gelatin/hydroxyapatite/bioactive glass (G/HA/CEL2) films with different compositions(100:0:0 (G1); 30:70:0 (G2); 30:0:70 (G3); 30:35:35 (G4) (%, w/w/w)) were prepared as scaffold mate-rials for tissue-engineering applications, particularly in the field of bone repair. A bioactive glass with 45%SiO2, 3% P2O5, 26% CaO, 7% MgO, 15% Na2O and 4% K2O molar composition was selected (CEL2).Genipin was used as a cross-linker for the gelatin component. Samples were characterized in terms of theirbioactivity, thermal properties, mechanical behaviour and cell compatibility. After only 3 days of incubationin simulated body fluid (SBF) at 37◦C, calcium phosphate crystals precipitated on G3 and G4 surfaces,due to the high CEL2 bioactivity. Cross-linking increased the thermal stability of the gelatine componentas indicated by thermal analysis (denaturation temperature was 92.3◦C and 97.6◦C for not cross-linked andcross-linked gelatin, respectively). Furthermore, tensile modulus of samples increased with increasing theinorganic phase amount (from 4.72 ± 0.23 MPa for G1 to 6.46 ± 0.05 MPa for G4). The adhesion andproliferation of human primary osteoblasts on composite films was evaluated. Cell viability was high withrespect to the control for all samples and the presence of hydroxyapatite exerted an important role in theability of mineralization.© Koninklijke Brill NV, Leiden, 2010

KeywordsBioactive glass, composite, gelatin, hydroxyapatite, tissue engineering

* To whom correspondence should be addressed. Tel.: (39-11) 564-6969; Fax: (39-11) 564-6999; e-mail:[email protected]

© Koninklijke Brill NV, Leiden, 2010 DOI:10.1163/092050609X12481751806213

1208 P. Gentile et al. / Journal of Biomaterials Science 21 (2010) 1207–1226

1. Introduction

Every year, millions of patients are suffering from bone defects arising from trauma,tumours or bone diseases. Thus, the requirement for new bone to replace or re-store the function of traumatized, damaged, or lost bone is a major clinical andsocio-economic need. Successful bone formation strategies have to yield bone withadequate functional and mechanical properties. Bone tissue engineering has beenused as the alternative strategy to regenerate bone [1], developing a variety of sys-tems able to mimic the specifically organized nanoscale structure of the tissue.

Since Nyman et al. successfully cured periodontal diseases using the barriermembrane technique (BMT) in 1982, guided bone regeneration (GBR) applica-tions have been successively introduced for clinic use [2]. Currently, bioresorbablemembranes have been highlighted because of their absence of a second surgery,and have come to be widely used in barrier membranes [3, 4]. Liao et al. [5] haveprepared collagen membranes with excellent cell affinity and biocompatibility toregenerate tissues; however, the mechanical strength of the membranes was poor.Furthermore, degradation rate of collagen membranes did not match normal tis-sue healing process. Then, the researchers looked for other methods to meet thedemands of degraded GTR membrane both in mechanical properties and biocom-patibility. Among the different solutions suggested by researchers [6, 7], compositesbased on apatite crystals and natural polymers have received increasing attentiondue to their ability to preserve the structural and biological functions of the dam-aged hard tissues in a more efficient and biomimetic way [8]. These compositeshave been formulated to have adequate properties for applications in the field ofbone repair, such as bioactivity, osteoconduction, osteoinduction and biocompati-bility [9].

Gelatin is a cheap, commercially available biomaterial that has gained interest inbiomedical engineering, mainly because of its biodegradability. Gelatin is obtainedby thermal denaturation or physical and chemical degradation of collagen, the mostwidespread protein in the body occurring in most connective tissues as skin, tendonand bone [10]. With respect to collagen, gelatin does not express antigenicity inphysiological conditions, it is completely resorbable in vivo and its physicochemi-cal properties can be suitably modulated [11]. Since biodegradable polymers maybe rapidly in vivo reabsorbed, cross-linking strategies have been applied to prolonggelatin resistance in vivo and to improve its mechanical properties [12]. Much in-terest has been recently addressed towards enzymatically- [13] or naturally-derivedcross-linking agents [14], with a low toxicity. Genipin, the aglycone of geniposide(an iridoid glucoside isolated from the fruits of Genipa americana and Gardeniajasminoides Ellis) has been shown to possess cross-linking activity towards aminocontaining materials [15]. Genipin has been recently used as a cross-linking agentof gelatin microcapsules for drug delivery [16] and of gelatin conduits for periph-eral nerve regeneration [17]. This natural cross-linking reagent has been reported tobe less toxic than glutaraldehyde and ideal for clinical use.

P. Gentile et al. / Journal of Biomaterials Science 21 (2010) 1207–1226 1209

Several composites of ceramic powder and degradable natural (gelatin, colla-gen) or synthetic polymers (poly(lactic acid) and poly(lactide-co-glycolide)) havebeen produced and investigated [18], with the aim to confer high bioactivity andadequate mechanical properties to the scaffolds. Hydroxyapatite (HA), tricalciumphosphate (TCP) and bioactive glasses are ceramics that have been widely used andinvestigated as bone substitutes in dense and porous form [19]. When implanted invivo, these materials are non-toxic, antigenically inactive, do not induce cancer andbond directly to bone without any intervening connective tissue layer. Recent re-ports on ectopic bone formation (osteoinduction or material-induced osteogenesis)of calcium phosphate biomaterials showed that osteoinduction might be an intrinsicproperty of calcium phosphate biomaterials [20, 21].

The contribution of this work consists of the production and characterisationof gelatin-based films containing synthetic HA and a bioactive glass (CEL2) withsimilar composition to normal human bone.

HA has been used extensively for bone regeneration. HA has the ability to inducemesenchymal stem cells differentiation towards osteoblasts [22], whereas nano-sized HA has been found to improve osteoconductivity due to its similarity withmorphology of bone minerals [23, 24]. The addition of HA in the gelatin/HA com-posites has been found to improve the mechanical properties of scaffolds and theactivity and viability of rat osteoprogenitor cells cultured on them [25]. On theother hand, in recent studies, bioactive glasses and glass-ceramics have been inves-tigated for tissue-engineering applications in bone repair and have been successfullyused in some clinical applications [26]. Distinctive properties of bioactive glassesare their ability to induce the precipitation of a microcrystalline HA layer whenput in contact with body fluids and in aqueous solutions containing calcium andphosphate ions, as well as to bond directly to bone. The addition of glasses ofthe SiO2–Na2O–CaO–P2O5 system has led to a greater bioactivity than HA, asreported in several studies [27]. The potential bioactivity of glass/HA compositecoatings on titanium alloy and Ti6Al4V has been investigated in a recent study [28]and it has been found that pure HA coating and the pure bioactive glass coatinglost about 40% in its bond strength when exposed to a high humidity environ-ment. On the other hand, the HA/bioactive glass composite coating performed betterthan HA and bioactive glass separately. Physico-chemical and mechanical prop-erties of HA/glass porous scaffolds for tissue engineering applications have beeninvestigated by Roohani Esfahani et al. [29]. Zhitomirsky et al. have characterizedbioactive glass/HA/chitosan and bioactive glass/HA/alginate coatings on varioussubstrates [30].

This work is focused on the preparation and characterizations of gelatin/HA/bioactive glass composites with various weight ratios between blend compo-nents, with the aim to select an optimal composition of the composites for bonerepair.


2. Materials and Methods

2.1. Materials

Type A gelatin (CAS No. G2500-100G) from porcine skin was supplied by Sigma(Milan, Italy). HA powder (submicrometric dimension, with clusters of a few mi-crometers, Fig. 1a), produced by a precipitation method, was supplied by Istituto diRicerca Tecnologiche per la Ceramica (IRTEC-CNR, Faenza, Italy) and a bioactiveglass (CEL2, particle size <30 µm, Fig. 1b) was prepared according to the pro-cedure reported by Vitale-Brovarone et al. [31]. The chosen glass belongs to theSiO2–P2O5–CaO–MgO–Na2O–K2O system and has the following molar compo-sition: 45% SiO2,3% P2O5,26% CaO, 7% MgO, 15% Na2O, 4% K2O. Genipin(GP) was purchased from Challenge Bioproducts (Taiwan). All solvents used wereof analytical grade and used without further purification.

Figure 1. Morphology of hydroxyapatite clusters (a) and CEL2 particles (b).


2.2. Methods

2.2.1. Preparation of Cross-linked G/HA/CEL2 FilmsComposites containing gelatin, hydroxyapatite and CEL2 glass were produced ac-cording to the following procedure. Gelatin (1 g) was dissolved in demineralisedwater (20 ml) at 50◦C by stirring for 30 min, obtaining 5% (w/v) solution. Then,HA and CEL2 were added to the gelatin solution in order to obtain G/HA/CEL2composites with various weight ratios between the components: 100:0:0; 30:70:0;30:0:70; 30:35:35 (%, w/w/w), respectively, coded as G1, G2, G3 and G4.

For cross-linked samples, GP was added to G/HA/CEL2 solution at definedweight percentage (2.5%, w/w with respect to the gelatin amount). Each mixturewas kept at 50◦C under stirring until a gel started to form. The gel was spreadon Petri dishes and then left drying at room temperature for 48 h. After drying,samples were washed several times in demineralised water to remove GP residues.Cross-linked composite samples were coded as G1, G2, G3 and G4.

2.2.2. Bioactivity AvaluationIn order to study the bioactivity of samples, cross-linked films (10 mm × 10 mmsize; thickness of around 60 µm for gelatin and 240 µm for the composites, asmeasured by means of a calibre) were soaked in 5 ml of a simulated body fluid(SBF) prepared according to the protocol described by Kokubo et al. [26], at 37◦Cand pH 7.4 for various time intervals (3, 7 and 14 days, solution was refreshed onceevery 2 days). SBF has a composition similar to human blood plasma and has beenextensively used for the in vitro bioactivity test [30]. At the end of each experiment,the specimens were removed from SBF and then abundantly rinsed with deionizedwater and dried for morphological analysis (SEM) and compositional examination(energy-dispersive spectroscopy; EDS).

2.2.3. Mechanical CharacterizationTensile tests of G/HA/CEL2 film samples in a wet state were measured us-ing a mechanical testing machine (MTS, QTest/10). Strip-shaped film samples(5 mm × 40 mm size and thickness of 60 µm for G1 and 240 µm for G2, G3 andG4, as measured by means of a calibre) were immersed in PBS for 30 min. Then,films were dried at the surface by gentle contact with a filter paper and immediatelystretched at a cross-head speed of 2 mm/min. In this study, five samples were evalu-ated for each composition. The Young modulus of each sample was calculated fromthe slope of the initial linear region of the stress–strain curve. The Young modulusof each composition was obtained as the average of the five measured values.

2.2.4. Differential Scanning Calorimetry (DSC)DSC experiments were carried out in a DSC 7 Perkin Elmer on specimens (5–10 mg) packed in aluminium pans. Non-isothermal scans were performed betweenroom temperature and 200◦C at a heating rate of 10◦C/min under nitrogen at-mosphere. Denaturation temperature and enthalpy were calculated as the temper-


ature of the maximum value of the denaturation endotherm and the peak area,respectively. Denaturation enthalpy was normalised with respect to gelatin content.

2.2.5. Cell CultureBefore cell seeding, material samples (10 mm × 10 mm size) were sterilized in70% ethyl alcohol (EtOH, Sigma) for 30 min, washed in PBS and incubated with anappropriate medium for 3 h. The medium was then discarded. Primary osteoblastswere grown from explants of human trabecular bone fragments from knee jointstaken at surgery (kindly provided by the Orthopedic Institute, “SS Antonio e BiagioHospital”, Alessandria, Italy). The osteoblasts were cultured in Iscove’s modifiedDulbecco’s medium supplemented with 10% fetal calf serum (Sigma), 2 mM L-glutamine (Sigma), penicillin (100 U/ml) and streptomycin (100 µg/ml) (Sigma)for 3 weeks as previously described [33]. Cells from up to three passages were usedfor all experiments. 2 × 104 cells/cm2 were seeded onto the materials to be testedand allowed to adhere and proliferate for 7 days, respectively.

2.2.5.1. Cell Proliferation (MTS Assay). The CellTiter 96® AQueous Assay(Promega, Italy) uses the tetrazolium compound (3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium, inner salt; MTS)and the electron coupling reagent, phenazine methosulfate (PMS). MTS is chem-ically reduced by cells into formazan, which is soluble in tissue culture medium.The measurement of the absorbance of the formazan can be carried out using 96-well microplates at 492 nm after 3 h of incubation at 37◦C. The assay measuresdehydrogenase enzyme activity found in metabolically active cells. Cells seeded oncell culture plate, adequately treated for cell adhesion, were used as controls. Themean and the standard deviations were obtained from three different experimentson the same specimen.

2.2.5.2. Mineralization Test. The alkaline phosphatase activity was measured us-ing a colorimetric assay with p-nitrophenyl phosphate (Sigma). After 7 days cul-ture, culture media from experimental groups were collected and p-nitrophenylphosphate substrate was added. The samples were incubated at 37◦C for 30 minin the dark. The reaction was stopped with 1 M NaOH (Sigma), and the absorbancewas read at 405 nm using a microplate reader. Cells grown in basic medium wereused as a negative control, instead cells grown in culture medium supplementedwith mineralizating factors (ascorbic acid and dexamethasone) has been used aspositive control. The mean and the standard deviation were obtained from threedifferent experiments.

2.2.6. Morphological and Compositional Characterization (SEM-EDS)Morphological analysis (SEM; Philips 525M) and compositional analysis (energy-dispersive spectroscopy, EDS, Philips EDS 9100) were performed on surfaces andfractured sections of all composite specimens. The samples were sputter-coatedwith silver prior to examination. Samples from cell culture tests were washed twicein 0.15 M cacodylate buffer (Sigma) and fixed for 30 min at 4◦C with Karnowsky


solution (2% paraformaldehyde and 2.5% glutaraldehyde in 0.15 M cacodylatebuffer, pH 7.2–7.4, Sigma), dehydrated in increasing EtOH (Sigma) concentrations,dried and silver-sputtered, as described earlier, before SEM examination.

2.2.7. Statistical MethodsAll quantitative data were presented as mean ± SD, unless noted otherwise. Sta-tistical analysis was carried out using single-factor analysis of variance (ANOVA).A value of P < 0.05 was considered statistically significant.

3. Results and Discussion

3.1. SEM Analysis

SEM analysis was performed on cross-linked G/HA/CEL2 composite films to eval-uate the effect of composition on sample morphology. Figure 2 shows SEM imagesof the upper (Fig. 2a–d) and lower surfaces (Fig. 2e–h) with the corresponding EDSspectra and SEM micrographs of the fractured sections (Fig. 2i–l) of cross-linkedG/HA/CEL2 films.

Figure 2a, e, i shows exemplary images of the typical morphology of GP-cross-linked gelatin films: upper (Fig. 2a) and lower (Fig. 2e) film surfaces were regularand smooth. The typical fibrillar structure of collagen was not detected, due to thedenaturation process of collagen to gelatin affecting the tertiary and quaternarystructure [34]. Although the calorimetric data (see Section 3.4) showed that notall the triple helix structure of native collagen was transformed into a random coilconformation, such a structure was not detected by SEM analysis. Gelatin films hada homogeneous thickness of 60–70 µm (Fig. 2i).

Figure 2b, f, j, c, g, k and d, h, l shows the upper and lower surfaces and thefractured section of G2, G3 and G4 films, respectively. For these compositions,a bi-layered structure was evidenced by SEM analysis and EDS spectra: the sectionimages (Fig. 2j–l) show a clear separation between a lower porous layer with arough surface (Fig. 2f–h) and an upper compact layer with smooth surface (Fig. 2b–d) with a relative thickness depending on composition.

EDS analysis showed that G2 samples consisted of a top gelatin layer (Fig. 2b)and a bottom layer mainly consisting of HA and a very low gelatin amount (Fig. 2f).A similar laminated structure has been documented by Bigi et al. for HA/G films[35]. HA was not water-soluble and simply dispersed inside the gelatin solution:the layered composition was, thus, the result of HA deposition on the bottom ofthe Petri dish during solvent casting, due to the high HA (3.16 g/cm3) density andthe slow gelatin gelling rate at room temperature. G2 films had a thickness of 250–260 µm (Fig. 2j), which was significantly higher than for G1 samples.

Similarly, for G3 films, two adjacent layers was distinguished (Fig. 2k), mainlyconsisting of a top homogeneous G layer (Fig. 2c) and a bottom CEL2 layer(Fig. 2g) as shown by EDS analysis. However, for G3 samples, the thickness ofthe top gelatin layer was lower than for G2 samples. This behaviour was attributed


Figure 2. SEM micrographs and EDS spectra of G/HA/CEL2 composite films. (a–d) Upper surfaceof (a) G1, (b) G2, (c) G3, (d) G4. (e–h) Lower surface of (e) G1, (f) G2, (g) G3, (h) G4. (i–l) SEMmicrographs of fractured section of (i) G1, (j) G2, (k) G3, (l) G4. Bar = 100 µm.

to both the well-known CEL2 surface reactivity, leading to a better mixing withgelatin, and the lower density of CEL2 (2.6 g/cm3) as compared to HA, that de-creased the amount of decanted bioactive glass during solvent casting. G3 films hada thickness of 230–240 µm (Fig. 2k).

A layered structure of the section was also found for G4 samples (Fig. 2l and d,h). In this case, thickness of the gelatin top-layer was intermediate between that ofG2 and G3 samples (the trend was G2 > G4 > G3), due to the intermediate contentof HA and CEL2 in G4 samples. G4 composites had a thickness of 240–250 µm(Fig. 2l).

SEM analysis suggested that the mixing of gelatin with HA leads to samples witha neat phase-separated structure. On the other hand, composites based on CEL2 andgelatin showed some degree of mixing, as suggested by the inclusion of a highergelatin amount inside the inorganic phase-based layer as compared to samples con-taining HA. Finally, samples containing HA/CEL2 as an inorganic phase showedan intermediate behavior. Similarly, the thickness of the composite films was a re-


Figure 2. (Continued.)

sult of the different mixing degree between the phases: the higher mixing degreebetween gelatin and the inorganic phase(s) led to an increased compaction and areduced thickness.

3.2. Bioactivity Evaluation

The aim of bioactivity evaluation is to investigate the formation of a bone-like car-bonated apatite layer on the surface of a substrate after soaking in SBF. Figure 3shows the upper (Fig. 3a–d) and lower (Fig. 3e–h) surfaces of composite samplesafter 3 days incubation in SBF. For G1 and G2 apatite precipitation was not evidentafter 3 days soaking in SBF, as suggested by SEM and EDS analyses (Fig. 3a, b ande, f). On the other hand, globular shaped agglomerates were found on the upper andlower surfaces of G3: EDS analysis indicated the presence of calcium phosphateson these surfaces (Fig. 3c, g). Bioactive glasses have unique properties if comparedto ceramic materials such as HA and β-TCP, and their composition can be tunedto obtain a material with a tailored reactivity in the human body, ranging from aslightly bioactive behaviour to a completely bioresorbable material [36]. CEL2 is ahighly bioactive glass and the ability of CEL2 in the form of macroporous scaffoldsto induce the precipitation of HA has been previously reported [31, 37]. Similarly,


Figure 3. SEM micrographs and EDS spectra of G/HA/CEL2 composite films after 3 days of immer-sion in SBF. (a–d) Upper surface of (a) G1, (b) G2, (c) G3, (d) G4. (e–h) Lower surface of (e) G1,(f) G2, (g) G3, (h) G4. Bar = 30 µm.

after 3 days soaking in SBF, calcium and phosphorus based crystals deposited onthe upper surface of G4 samples (Fig. 3d), which contained a lower CEL2 amountas compared to G3 samples (Fig. 3c).

SEM images and EDS spectra reported in Fig. 4 show the upper (Fig. 4a–d)and lower (Fig. 4e–h) surfaces of composite samples, after 1 week incubation in


Figure 4. SEM micrographs and EDS spectra of G/HA/CEL2 composite films after 7 days of immer-sion in SBF. (a–d) Upper surface of (a) G1, (b) G2, (c) G3, (d) G4. (e–h) Lower surface of (e) G1,(f) G2, (g) G3, (h) G4. Bar = 100 µm.

SBF. After 1 week, G2 and G4 upper surfaces were homogenously covered bya thin white globular shaped agglomerate-layer (Fig. 4b, d), rich in calcium andphosphorous.

Bioactivity of G3 was stronger than that of G2 and G4 samples, and it was com-pletely covered by calcium phosphates after one week immersion in SBF (Fig. 4c).


Moreover, for G3, the apatite layer formed on both the upper and the lower filmsurface, due to the higher bioactivity of CEL2 particle than HA agglomerates andprobably to the better CEL2 particles dispersion in the film [33].

SEM micrographs and EDS spectra reported in Fig. 5 show the upper (Fig. 5a–d)and lower (Fig. 5e–h) surfaces of composite samples, after 2 weeks incubation in

Figure 5. SEM micrographs and EDS spectra of G/HA/CEL2 composite films after 14 days of im-mersion in SB. (a–d) Upper surface of (a) G1, (b) G2, (c) G3, (d) G4. (e–h) Lower surface of (e) G1,(f) G2, (g) G3, (h) G4. Bar = 100 µm.


SBF. After 2 weeks of immersion in SBF (Fig. 5b–d and f–h), for each compos-ite sample, the calcium and phosphorus amounts as measured by EDS analysis onthe upper and lower surfaces, increased significantly, as compared to the previousanalysed time intervals. The calculated Ca/P ratio for each sample was about 1.8,which is close to the Ca/P ratio of the human bone, which ranges between 1.6 and1.7 depending on the age [32]. On the other hand, pure gelatine samples did notinduce the precipitation of calcium and phosphorus containing crystals at any testtime (Fig. 5a, e), suggesting that the inorganic phase had a key role in determiningthe bioactivity of composites [20].

It is well known that bioactive glass has a higher bioactivity than HA, whichmakes it a more suitable material for applications in bone reconstruction [38]. Ithas been shown, for example, that there is a faster bone formation in one weekin the presence of bioactive glass than that is formed when HA or other calciumphosphate ceramic particulates are placed in the same type of defect [39]. Moreover,bioactive glass, being a class-A bioactive material (as opposed to HA, which is classB), has shown strong bond also to soft tissues [40]. The application potential of thebioactive glass-containing composites fabricated here, therefore, could include bothhard and soft tissue regeneration and repair.

Figure 6 reports the pH value of SBF media measured after sample immersion atfixed time intervals (1, 3, 5, 7, 10 and 14 days). For G1 and G2 samples, pH slightlyincreased in the first days of incubation in SBF and then settled approximately toa constant value of 7.55–7.65. For samples containing CEL2, pH of SBF mediumcontinued to increase until the second day of sample incubation in SBF, then pHdecreased on the third day and settled at an approximately constant value (7.7–7.8).

The measurement of pH variations was crucial for samples containing CEL2(G3 and G4), as they displayed the highest pH at each soaking time, due to the

Figure 6. pH value of SBF solution for GP-cross-linked G/HA/CEL2 film samples.


Figure 7. Young’s modulus of G/HA/CEL2 composites.

ion-leaching phenomena of CEL2 glass [31]. The evaluation of pH was studiedwith the aim to determine the more favourable sample pre-treatment times in SBFbefore performing biological tests, to avoid the release of ions from the substratesduring cell culturing which may lead to pH changes and consequent cell death.Our results suggested that during the first week of soaking in SBF, the rate of ionrelease was quite high and induced measurable pH variations, which are knownto be detrimental for cell survival: changes in extracellular pH are likely to causeapoptosis or necrosis, as the potential across the cell membrane can be altered.However, after 1 week of soaking in SBF, the pH value was 7.7–7.8 for samplescontaining CEL2, which is suitable for in vitro cell testing [41].

3.3. Mechanical Characterization

Tensile tests were performed on wet samples, previously soaked in PBS for 30 min.Figure 7 shows the behaviour of the elastic modulus of film samples as a function ofG/HA/CEL2 weight ratios and GP amounts. As a control, an uncross-linked gelatinsample was tested. Genipin cross-linking improved the mechanical properties ofgelatin films (0.89 ± 0.08 MPa for un-cross-linked gelatin vs 4.72 ± 0.23 MPa forG1).

On the other hand, Young’s moduli were not significantly different for G3 andG4 composite films and their values were slightly lower than G2 samples but higherthan G1 samples (5.62 ± 0.19 MPa for G3, 6.46 ± 0.05 MPa for G4).

It is widely accepted that tensile strength for inorganic–organic hybrid materialscan be viewed as an indication of the interfacial bonding strength between the min-eral and organic components [42]. In addition, the proper stress transfer occurringbetween the mineral phase and the organic matrix has a major effect on the mechan-ical properties of the composite materials [43]. Regarding the present context, thechemical interaction and the intimate adhesion between inorganic powders and theorganic matrix, as well as the superior stiffness of the organic phase were the reasonfor the enhanced tensile strength of composites as compared to gelatin samples. Assuggested by Bigi et al. [35], the presence of the apatitic crystals in the compositefilms induced an increase of stiffness of HA/gelatin, as suggested by the increase ofthe Young’s modulus and the consequent decrease of elongation-at-break.


Figure 8. DSC thermograms of (a) un-cross-linked and cross-linked gelatin films, and (b) com-posite films. This figure is published in colour in the online edition that can be accessed viahttp://www.brill.nl/jbs

3.4. DSC

DSC analysis was performed to analyse the thermal behaviour of composite mate-rials as a function of composition and the influence of GP-cross-linking on gelatinethermal properties. Figure 8a shows a DSC thermogram of a wet un-cross-linkedgelatin sample, which exhibits an endothermic peak centred at about 90.2◦C, as-sociated with the helix–coil transition of gelatin, with a denaturation enthalpy ofabout 30 J/g.


Table 1.Denaturation enthalpies (�Hd) and denaturation tem-peratures of G/HA/CEL2 composites

�Hd (J/g) Td (◦C)

G/HA/CEL2_2.5GPG1 23.3 97.6G2 21.3 99.6G3 9.3 98.6G4 9.0 99.0

G/HA/CEL2_0GPG1 30.3 92.3

Cross-linking increased the thermal stability of gelatin helices, as shown bythe shift of the denaturation temperature to higher values (97.6◦C for G1). Cross-linking generally induced a decrease in the denaturation enthalpy, which was as-cribed both to a reduction of hydrogen bonds, which break endothermically, and toa simultaneous increase in the extent of covalent cross-links, which break exother-mically [44] (30.3 J/g for uncross-linked gelatin and 23.3 J/g for G1). The val-ues of denaturation temperature (Td) and denaturation enthalpy (Hd) obtained forG/HA/CEL2 composite films are reported in Table 1, whereas the relative thermo-grams are shown in Fig. 8b.

The denaturation temperature of G/HA/CEL2 composites with different weightratios between G/HA/CEL2 was slightly increased with respect to pure cross-linkedgelatin film (99.6◦C for G2, 98.6◦C for G3 and 97.6◦C for G1). It is noteworthy thatcomposites containing CEL2 (G3 and G4) showed low denaturation enthalpy values(9.3 J/g for G3 and 9.0 J/g for G4), probably due a reduction of the helical structureas a consequence of the strong interactions between bioactive glass and gelatin, aswell as the fine distribution of the components (as shown by SEM analysis (Sec-tion 3.1)).

3.5. Cell-Culture Tests

The adhesion and proliferation of human primary osteoblasts on composite cross-linked films was evaluated to determine the effect of composition on the cellresponse, as for each clinical case, this parameter should be tailored to the require-ments of the implantation site. Figure 9 depicts the morphology of primary os-teoblast cells adhering to composite films. SEM micrographs showed that primaryosteoblasts had a high affinity for the samples, displaying the typical osteoblast-likemorphology, with a stretched elongated shape on G2 samples (data not shown), anda spread morphology on scaffolds with CEL2 (Fig. 9a, b). It seemed that osteoblastcells were located under the surface of composite films, suggesting the material had


Figure 9. SEM micrographs of primary osteoblast cells seeded onto (a) G3 and (b) G4. Bar = 20 µm.

resorbed partially and, as a result, the osteoblasts had buried themselves during thisprocess, as also reported by Zhao et al. [45]. As a consequence of that, cells mayprogressively come into contact with the underlying inorganic layer.

As shown in Fig. 10a, after 7 days of culture, human primary osteoblasts viabil-ity evidenced no statistically differences between the control and the G1, G3 andG4 samples. The optical density had an average value of roughly 0.06 for all mate-rials. Only G2 showed a higher viability value indicating an increased proliferationrate stimulated by the higher HA amount. However, all samples did not negativelyinterfere with cell viability as compared to the control.

Figure 10b shows the results of ALP activity tests. Alkaline phosphatase activ-ity increased when in presence of adequate biochemical stimuli (positive control)and not significantly when in absence (negative control). For the composites, theinhibition of alkaline phosphatase activity was particularly evident for G1, with sta-tistical relevance, and G4 with a high standard deviation which does not confirm thestatistic relevance. G3 showed an increased activity as well as the positive control.G2 seemed to strongly enhance the enzymatic activity, having a statistical differ-ence with the negative control and showing a value higher than the positive control.Thus, only in this case the material was able to stimulate the physiological activityof osteoblasts significantly.


Figure 10. MTS test and ALP activity in primary osteoblast culture on composite films with varyingcomposition (∗P < 0.05).

4. Conclusions

This work focused on the realisation and characterisation of biomimetic compos-ites based on 30 wt% gelatin and 70 wt% bioactive ceramics (HA/CEL2, in threedifferent relative compositions), to be used as scaffolds for bone repair. Genipincross-linking of the gelatin component was performed, resulting in an improvementof thermal and water stability as well as the mechanical properties. Composites con-taining CEL2 were particularly interesting due to their pronounced bioactivity andexpected consequent bone-bonding ability during in vivo trials. In fact, they areexpected to react with physiological fluids, forming hydroxyapatite layers on thefilm surface containing inorganic phase and creating strong bonds to hard and softtissues through cellular activity. Drawbacks of CEL2 include: (i) low mechanical


properties, with consequent poor mechanical performances of composites based onCEL2, and (ii) ion leaching phenomena in a water media, resulting in pH varia-tions causing cell death in some cases. Mixing CEL2 with HA (or possibly anotherbioactive ceramic) allowed to overcome the first drawback. On the other hand, a pre-incubation time in SBF of seven days was found to avoid pH changes of the culturemedium during biological tests. The obtained composites represent promising can-didates for future trials in the field of bone regeneration (as suggested by MTSand ALP tests), in which a laminated structure is interesting for in vivo applica-tions, where the inorganic layer may improve the scaffold adhesion to bone tissue.Additional work is in progress, with the aim to realise scaffolds with the same com-positions of the studied film-samples, but showing an interconnected pore network.

Acknowledgements

The authors wish to thank the European Commission for financial supportthrough STReP-project “Photozyme nanoparticle applications for water purifica-tion, textile finishing, photodynamic biomineralization and biomaterial coating”-PHOTONANOTECH/NMP-4-CT-2007-033168. The authors also thank Mr. MauroRaimondi for his technical support during SEM analysis.

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Composite Films of Gelatin and Hydroxyapatite/Bioactive Glass for Tissue-Engineering Applications

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