Article pubs.acs.org/Biomac
Effect of Cross-Linking Methods on Structure and Properties of Poly(ε-caprolactone) Stabilized Hydrogels Containing Biopolymers Geta David,† Mariana Cristea,‡ Ciprian Balhui,§ Daniel Timpu,‡ Florica Doroftei,‡ and Bogdan C. Simionescu*,†,‡ †
Department of Natural and Synthetic Polymers, “Gh. Asachi” Technical University of Iasi, 700050, Romania “Petru Poni” Institute of Macromolecular Chemistry, Iasi, 700487, Romania § ASIL Cosmetics, Galati-Tg.Bujor, Romania ‡
ABSTRACT: Different dense and porous biodegradable matrices based on solely atelocollagen, or with different atelocollagen and hyaluronic acid derivative ratios, were obtained by varying feeding formulations, crosslinking reaction parameters, and preparative protocols. The compositions and methods for forming hydrogels through a combination of physical and chemical cross-linking processes are provided. The chemical cross-linking was mainly mediated by a synthetic component, a poly(ε-caprolactone) reactive derivative, aiming the development of new hybrid hydrogels with tailored characteristics by an appropriate use of the advantages offered by the included natural and synthetic components and the selection of the preparative procedure. The structure and morphology of the 3D hybrid materials were comparatively investigated by means of Fourier-transform infrared spectroscopy (FTIR), differential scanning calorimetry (DSC), Xray diffraction (XRD), and environmental scanning electron microscopy (ESEM). FTIR and XRD analysis showed no signs of collagen denaturation during the formation of 3D structures. The influence of various factors, such as the chemical composition of the resulted hydrogels and their morphology, on water uptake and water vapor sorption, mechanical behavior, as well as on in vitro degradation characteristics, was systematically investigated. The experimental results point on the advantage offered by the high and modular physicochemical stability of the ternary hydrogels cross-linked by combined approaches. All newly developed materials show no hemolytic effect, which recommends them for potential biomedical applications.
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INTRODUCTION The ability of hydrogels to mimic body tissues recommends them as promising biomaterials in a large range of applications in pharmaceutical, medical and cosmetics industry.1−3 Their drawbacks, as the poor strength characteristics, may be diminished or avoided by a careful selection of the appropriate composition and preparative methodology, and different strategies, including copolymerization, cross-linking, and formation of interpenetrating polymer networks or composite structures, have been proposed. Composites of synthetic and natural polymers benefit from the wide range of physicochemical properties and processing techniques related to synthetic polymers and the bioactivity of the natural ones. As an example, collagen/poly(ε-caprolactone) (Col/PCL) biocomposite membranes were prepared and investigated in the search for an ideal support in tissue engineered skin replacements, with the goal to overcome the limitations of natural grafts for treatment of burn injuries and chronic wounds.4 Conventional autologous skin transplantation presents disadvantages such as the need for harvesting and limited supply, while dermis-containing allografts carry a risk of immune rejection and disease transfer. © 2012 American Chemical Society
The collagen-glycosaminoglycan (CG) system was found to provide a substrate for epidermis repair, substrate with improved stability, and high mechanical strength relative to pure collagen.5 Various CG systems were developed and commercialized as skin substitutes.6,7 However, marketed versions of cross-linked CG porous scaffolds also have their limitations, for example, risk of wound infection.8 In this context, considering the continuous increasing interest for the development of improved materials for wound healing, the present study is aiming to provide new CG systems with tailored characteristics. This objective is intended to be achieved through the correlation of the properties of the developed materials with their formulation and the applied preparative method. A ternary composition, comprising collagen, hyaluronic acid (HA), and poly(εcaprolactone), was proposed. Received: March 23, 2012 Revised: May 18, 2012 Published: June 13, 2012 2263
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Cross-linking agents play an important part in the physical properties of collagen based biomaterials. According to the most recent data, double-cross-linked networks are characterized by improved performances in use together with suitable structural hierarchies.9,10 As a consequence, and considering also our earlier work, combined cross-linking systems were applied, a PCL reactive derivative was used as a long-range cross-linker, while UV irradiation, silanol cross-linkers, or a nontoxic cross-linking system based on water-soluble carbodiimide and N-hydroxysuccimide (EDC/NHS) were added to complete its effect on the physicochemical stability of the developed material.11 The synthetic biodegradable semicrystalline polyester PCL is a material commonly used in combination with natural polymers to control the biodegradation rate and mechanical properties in biomaterials.12 Collagen, a primary component of extracellular matrices (ECM), is an accessible, biocompatible, completely in vivo resorbable protein, able to support cell growth and differentiation, presenting negligible immuno-reactivity and properties that can be adapted to successfully meet a range of different clinical applications. It is already applied in various forms (fibers, gel, foam), accompanied by other materials, in the treatment of wounds and burns.13 In this study, purified atelocollagen I was used.14 Hyaluronic acid, a nonsulfated glycosaminoglycan (GAG), is another essential component of the ECM. Due to its high biocompatibility, biological activity, and reach chemistry, HA as so or its derivatives have been clinically used as products of various physical forms, viscoelastic solutions, soft or stiff hydrogels, electrospun fibers, sponges, and nanoparticulate fluids, for some decades.15 It gained popularity mainly as a good base for a soft tissue augmentation material, being used from the beginning as a bioactive ingredient in skin care products, in respect to its important role in the maintenance of skin elasticity, suppleness, and tonicity. Recently, HA became an important component in the design of new biomaterials for regenerative medicine, its market position competing with that of collagen. In this context, dimethylsilanediol hyaluronate (DMSHA) was chosen as a component in the present study. Silanols are known as bioactive silicon derivatives with antiinflammatory, cyto-stimulating, and cyto-protection properties, with a role in restructuring and metabolic modulation of connective tissue.16 Such compounds were studied and included in different forms as a passive or active component in different topical patented or commercial dermatological or cosmetic compositions for treating symptoms of skin aging, skin defects, or wounds, in formulations for personal care products,17,18 hydrophilic adhesives for medical skin coverings,19 and so on. The commercial product used in this study (DSH−CN) was subjected to in vitro tolerance tests on both cell culture (fibroblasts) and reconstituted epidermis, the tests evidencing that this HA derivative is neither toxic nor irritant and has a cyto-stimulating effect.17,20 The hybrid 3D materials prepared by two different strategies, to obtain dense or porous structures, were comparatively characterized and different structural, morphological, or behavior aspects were discussed in terms of the peculiar relation between materials design, their formulation and their performances, in order to evaluate the possibility of in-use properties control.
Article
EXPERIMENTAL SECTION
Materials. Native type I collagen was isolated from bovine Achilles tendon and further used to prepare aqueous acidic solutions (pH ∼ 2) of purified type I atelocollagen (AteCol), according literature.14 Dimethylsilanediol hyaluronate (DMSHA) solution (commercial form, DSH−CN: aqueous solution with a pH of 5.5, containing 0.3% dimethylsilanediol/0.09% Si, and 0.3% hyaluronan, the sodium salt of hyaluronic acid of nonanimal origin, with Mw of 1.8−2.2 MDa) was supplied by EXSIMOL S.A.M. (Monaco). Poly(ε-caprolactone) glycol (Mn = 2000), 4,4′-methylenebis-(cyclohexyl isocyanate) (H12MDI), Triton X-100, poly(vinyl pyrrolidone) K15, 1-ethyl-3(3-dimethylaminopropyl)carbodiimide (EDC), N-hydroxysulfosuccinimide (NHS), and dimethylsulfoxide (DMSO) were purchased from Fluka (Germany). Poly(ε-caprolactone) diisocyanate (PCL-DI) was obtained according to an earlier report.11 All other solvents (ethanol, acetone) and reagents of analytical grade, commercially available, were used without further purification. Bidistilled water was used for formulations preparation. Membranes Preparation. Three spongy (procedure 1) and eight dense (procedure 2) membranes of different formulations were prepared, varying the ratios of AteCol, DMSHA and PCL-DI. These were subjected or not to short-range cross-linking, that is, (1) UV irradiation (Osram HBO 200 W super pressure mercury lamp, 14.6 W, luminous intensity Iv = 1100 cd) or (2) EDC/NHS cross-linking. Dense and porous un-cross-linked AteCol matrices, and another porous one containing only AteCol and DMSHA, were also prepared as references. The adopted samples code was CHxPy-z, n, where x is % DMSHA, y is % PCL-DI, z represents the irradiation duration (min), and n the preparative procedure (1 or 2). C, H, and P are the abbreviations used for AteCol, DMSHA, and PCL, respectively. Briefly, to obtain porous membranes (procedure 1), the appropriate amount of DSH−CN was blended dropwise with the AteCol dispersion (brought at a pH of 5.5 with NaOH 0.1 N), under stirring, for homogenization. For the formulations containing PCL, the calculated PCL-DI amounts in the appropriate volume of solvent (2:3 v/v acetone/DMSO, representing ∼1% relative to total dispersion volume) containing 4 wt % (relative to PCL-DI) Triton X-100 stabilizer, were added dropwise to the dispersions. The mixtures were homogenized 10 min by stirring and 5 min by sonication. After deaeration they were frozen (−20 °C) and lyophilized with a CHRIST freeze-dryer, Alpha 1−4 LSC type. The resulted sponges were irradiated according to the mentioned procedure. For comparison, a sample of CG lyophilized sponge was crosslinked by using the EDC/NHS system (CH1ENP2−30, 1) in ethanol/ water mixture (EDC/NHS/collagen carboxylic acid groups 10:10:1, ethanol mole concentration of 0.13), according to a recently developed alternative.21 After reacting for 14 h at 4 °C, the hydrogel was washed with saline (15 min), water (3 × 5 min), 0.1 m Na2HPO4 (pH 9.1, 30 min), phosphate buffer (3 × 30 min), water (3 × 5 min), citrate buffer (3 × 30 min). The resulted dispersion was mixed with an appropriate amount of PCL-DI in acetone−DMSO. The prepared mixtures were degassed to remove the bubbles, frozen at −20 °C and lyophilized. For dense matrices preparation (procedure 2), the AteCol (1.8 wt %) and AteCol/DMSHA/PCL-DI dispersions, prepared according to the above-described protocol, were cast on the dish (preliminary covered with a thin film of poly(vinyl pyrrolidone) as a demulant) after deaeration and then dehydrated by slow drying in an desiccator under vacuum until a film was formed. The dried film was or was not subjected to UV irradiation. Characterization. The FTIR spectra were obtained on a Vertex 70 (Bruker) spectrophotometer. The denaturation temperature (Td), indicative for the degree of matrix cross-linking, was measured using differential scanning calorimetry (DSC) with a Mettler 851 system. Samples were heated from −20 to 140 °C at a heating rate of 3 °C/min, in nitrogen atmosphere, with an empty aluminum pan as the reference. The denaturation temperature was determined as the peak value of the corresponding endothermic phenomena. Thermogravimetric analysis 2264
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Scheme 1. Schematic Representation of Atelocollagen−Dimethylsilandiol Hyaluronate−Poly(ε-caprolactone) Ternary Hydrogels Formation
and the corresponding moisture vapor adsorption rate (×10−3 %/s) were evaluated. To obtain the relative degradation rates of the samples with different compositions and morphologies, specimens were incubated at 37 °C in 0.1 N NaOH or phosphate buffer solution (PBS) for predetermined periods of time, then washed with water and freezedried. Mass loss was calculated by comparing the initial mass (W0) with that measured at a given time point (Wt), as shown in eq 3. For every specimen three individual experiments were performed for the degradation test.
(TGA) was performed under nitrogen stream (ramp rate of 10 °C min−1). The porosity values of the composite sponges were measured by liquid displacement, according to a published method.22 Ethyl alcohol was used as the displacement liquid. The surface and cross-section morphologies of the scaffolds were observed directly by a scanning electron microscope (SEM - Quanta 200 instrument, working in low vacuum mode) without sputter coating by conducting matter. The pore dimensions were estimated from SEM microphotographs by the manual measurement of pore zones with arbitrary shapes. Wide-angle X-ray diffraction (WAXD) investigations were performed using a D8 Advance (Bruker) diffractometer with an integrated detector with scintillation and a Cu Kα source with a wavelength of λ = 1.54 Å, at a scanning rate of 0.01° s−1, in the 2θ range of 3.0 to 40°. The flexural storage modulus (E′), loss modulus (E″), and the dissipation factor (tan δ) were determined via dynamic mechanical analysis (DMA) using a Pyris Diamond type instrument (PerkinElmer) over a temperature range of −100 to 300 °C, at a frequency of 1 Hz and a ramp of 2 °C min−1. Tensile tests were performed on approximately 10 × 9.5 × 0.2 mm thin film samples, at a constant displacement rate of 50 mN min−1, on the same instrument. The swelling capacity studies were performed at room temperature by immersing the weighed samples of 2 × 2 × 0.3 cm in bidistilled water. At specified time intervals the samples were taken out of the water, blotted to remove surface water and weighed. The mass swelling ratio (SR) was calculated by the relation
SR(g g −1) = (Ws − Wd)/Wd
mass loss(%) = (W0 − Wt)/W0 × 100
The hemocompatibility properties of the prepared materials were evaluated spectrophotometrically by hemolysis tests, according to %lysis = (C − Cn)/(C p − Cn) × 100
(4)
where C is hemoglobin concentration in the sample, Cn is the concentration of hemoglobin in the negative control (plasma separated from blood), and Cp is the concentration of hemoglobin in the positive control (blood hemolysed with distilled water). The hemoglobin concentration was determined with the polychromatic formula23 C(mg/mL) = 1.65A415 − 0.93A380 − 0.73A470
(5)
where C is hemoglobin concentration in mg/mL and A415, A380, and A470 are the absorbances at 380, 415, and 470 nm, expressed in milliabsorbance units. For each hydrogel sample and the corresponding positive and negative control, fresh blood collected from a single healthy human volunteer into tubes containing dipotassium ethylenediaminetetraacetate was used.
(1)
where Wd is the weight of the dry sample and Ws is the weight of the swollen one. The equilibrium swelling ratio (ESR) is the ratio after equilibrium swelling. The experimental plot was obtained from averages of three samples. Water vapor sorption measurements were performed on a IGAsorp system (Hiden Analytical, Warrington, U.K.) on a relative humidity domain of RH = 0−90%, after samples preliminary drying in a nitrogen flow (250 mL/min, 25 °C). The dynamic sorption capacity at RH = 90%, defined by eq 2 (%) = [(WRH = 90 − WRH = 0)/WRH = 0] × 100
(3)
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RESULTS AND DISCUSSION DMSHA and AteCol were combined in a range of compositions, in the presence or absence of PCL-DI as a long-range cross-linker, to produce matrices with various properties, depending on the used chemistry and the applied preparative procedure. The efficacy of the former developed PCL reactive derivative as a long-range cross-linker for
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Table 1. Composition of the Reaction Mixtures Used in the Preparation of Atelocollagen−Dimethylsylanediol Hyaluronate− Poly(ε-caprolactone) Matrices and Products Properties No.
code
1 2 3 4 5 6 7 8 9 10 11 12 13 14
AteCol, 1 AteCol, 2 CH1P5, 2 CH1P5−15, 2 CH10P5, 2 CH10P5−15, 2 CH1P30, 2 CH10P30−15, 2 CH10P30, 2 CH10P30−15, 2 CH10-P10−15, 1 CH1P2−30, 1 CH1ENP2−30, 1 CH1, 1
DMSHAa (wt %)
PCL-DIa (wt %)
1 1 10 10 1 1 10 10 10 1 1 1
5 5 5 5 30 30 30 30 10 2 2
cb (%) 0.8 1.3 1.0 1.0 0.8 0.8 1.0 1.0 0.8 0.8 0.8 0.3 0.8 0.3
UVc (min)
15 15 15 15 15 30 30
A1240/1450d 0.95 1.02 0.95 0.90 0.95 0.80 1.08 1.17 1.04 1.02 0.90 0.89 0.93 1.04
a Relative to AteCol, by weight. bGlobal concentration in final dispersion. cUV irradiation (Osram HBO 200 W high pressury mercury lamp, 14.6 W, Iv = 1100 cd). dFrom FT-IR spectra.
collagen-based hydrogels was demonstrated in an earlier report.11 According to preliminary data on cross-linked collagen sponges, a duration of 15 or 30 min UV irradiation was employed as a short-range cross-linking clean alternative, because it allows the temporal and spatial control and enhances enzymatic resistance without toxic side effects. For comparison, a sample was also submitted to short-range cross-linking by EDC/NHS system. The long-range and short-range bonds led to the formation of hydrogels with a hybrid network structure. Even if mixtures of collagen and PCL or collagen− glycosaminoglycan were previously widely studied,4−7 to our knowledge, no investigation was reported on ternary Col/ GAG/PCL hydrogels, as well as on PCL-based hybrid crosslinking system for the reinforcement of CG composites. The main reactions implied in the generation of stable ternary hybrid 3D systems are described in Scheme 1.11,15,24 Table 1 gives an overview on the composition and properties of the new prepared materials. Usually, matrices based on hyaluronic acid are sensitive to temperatures lower than 0 °C, these leading to high free HA (non-cross-linked HA) content. Contrary to the prejudice in the art, reducing the temperature of the mixture during processing to form a matrix (lyophilization procedure) did not result here in structural degradation, an effect also observed recently for 3D networks containing only hyaluronic acid (dermal filler), attributed to the double cross-linking, respectively, to the use of previously partially stabilized derivatives.15 The formed hydrogels can be designed to have structural strength, no matter the preparative route (porous or dense). Typical FT-IR spectra of the prepared matrices with different hyaluronan derivative and PCL content are shown in Figure 1. Distinctive absorption bands attributed to AteCol I are present at 3330, 3081 (associated with the stretching vibrations of N− H groups), 1650 (originated from CO stretching vibrations coupled to N−H bending vibrations), 1540 (N−H bending vibrations coupled to C−N stretching vibrations), and 1240 cm−1 (related to the C−N stretching ν(CN) and N−H in-plane deformation ν(NH) modes), denoted as amide bands A, B, I, II, and III, respectively. The bands occurring in the range from 2800 to 3050 cm−1 are associated with the C−H vibrational modes ν(CH2) and ν(CH3). The inclusion of hyaluronic acid
Figure 1. FTIR-ATR spectra of pure atelocollagen, dried DMSHA, and CHxPy matrices: (A) dense films; (B) porous matrices.
derivative and PCL in the formulation gives rise to the appearance of new signals situated at 1723 cm−1 (-O−CO-, ester stretching in PCL), 1261 cm−1 resulting from the primary O−H in-plane bending in polysaccharide, 1180 cm−1 (assigned to the stretching vibration of the ester group), and 1054/1070 and 1030/1025 cm−1 (stretching vibrations of C−O in polysaccharide ring, superposed with that of ester group in PCL). Generally, hydrogel formation did not imply modifications neither for the position nor the intensity of amide I band, situated at around 1650 cm−1 and related to triple helix conformation changes, indicating that the collagen triple helix is not affected.25 The maintenance of the native structure of collagen is confirmed by the values of the A1543/A1240 ratio, situated mostly around unity (Table 1). With the increase in hyaluronic acid level the amide A band is enlarging, by contribution of the broad band at 3483 cm−1 due to the hydroxyl stretching vibration of the polysaccharide, while the intensity of the bands situated at 1461, 1080, and 1020 cm−1 is increasing. The signal at 805 cm−1, assigned to Si−CH3, becomes visible. The increase of the PCL amount results in the increase of the intensity of the peaks situated at 1723, 1180, and 2266
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1030 cm−1 and a decrease of that of the amide A band, a consequence of the NH2 groups consuming by reaction with isocyanate groups from PCL-DI (Figure 1A,B). Amide band I is removing slightly from 1632 to about 1654 cm−1 only for the highest PCL content (30%), in both dense and porous materials. The absence of essential denaturation in collagen is also evidenced by the diffractograms of different samples (porous or dense matrices), shown in Figure 2, which are all presenting a
Figure 3. DSC scans of pure un-cross-linked AteCol and of different films with a ternary composition.
Figure 2. WAXD diffractograms for AteCol and DMSHA and some of the developed dense hybrid materials.
distinct peak centered at 2θ ∼ 7.7°, the characteristic equatorial peak of collagen. This is removing to lower or higher 2θ values, depending on formulation, i.e. the presence in the reaction medium and final composition of solely PCL or of both PCL and DMSHA. Most possible, this shifting is due to the insertion of PCL chains between fibrils and Col-HA biocomposite formation, respectively. The broad peak at 2θ = 15−28° indicates the position of the characteristic interchain spacing of the collagen triple helix. The presence of the signals of relatively narrow width at about 15−18°, and of lower order diffraction peaks around this position, indicates the maintenance of a certain degree of order, that is, crystallization for collagen fibrils. This is more evident for sample CH1P5−15, 2. This aspect is related to an increase of crystallinity with the introduction of a small amount of PCL in the formulation. By the contrary, all specific collagen peaks become broader with increasing hyaluronan and PCL content, suggesting a possible slight alteration of the collagen structure for the highest synthetic polyester amounts. In turn, narrow peaks, assigned to PCL microphase separation and some degree of crystallization, become visible at 22° in the diffractograms of the samples with a high percent of polyester. The cross-linking efficiency was evaluated by the investigation of samples thermal behavior and mechanical characteristics, completed by swelling ability and dynamic water vapor sorption, the modifications in molecular organization inducing variations in the material application properties.26 DSC and TG measurements were used to detect the temperature of denaturation, Td, for pure atelocollagen I and hybrid matrices based on it, and to assess their thermal stability (Figures 3 and 4 and Table 2). Td is associated to the transformation from crystalline triple helix to amorphous random coil in collagen, accompanied by the breakage of
Figure 4. Thermogravimetry (full lines) and derivative thermogravimetry (dotted lines) curves for (a) AteCol, 2; (b) CH10P5−15, 2; (c) CH1P30−15, 2.
Table 2. Thermal Characteristics of the Prepared Matrices code
Td (°C)
AteCol, 1 AteCol, 2 CH1P5, 2 CH1P5−15, 2 CH1P30−15, 2 CH10P5−15, 2 CH10P30, 2 CH10P30−15, 2 CH10P10−15, 1 CH1, 1 CH1P2−30, 1 CH1ENP2−30, 1
70.8 74.6 82.3 84.6 91.0 85.9 81.8 81.5 89.7 92.8 97.4 96.8
Tm (°C)
48.0 50.3 50.3 49.6 39.5
inter- and intramolecular hydrogen bonds on heating and the release of loosely bound water. Thus, the temperature of thermal denaturation strongly depends on the water content in collagen and its degree of cross-linking between the chains.27 A comparison of the obtained data shows that the addition of PCL and DMSHA causes an increase of the denaturation temperature from 70.78 °C to values ranging from 81.5 to 97.44 °C (Figure 3, Table 2). These facts indicate the 2267
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occurrence of the cross-linking process. The presence of HA, which also acts as a cross-linker during blending with collagen to give a composite, enhances the cross-linking effect with the consequent improvement of thermal stability and gel stiffness. Further increase of PCL content for dense samples with high HA amount yields lower Td values (Table 2), most probably due to the incomplete consumption of excessive PCL-DI. The excessive PCL presence may also explain the appearance in DSC thermograms of the characteristic melting point (Tm) of PCL at around 50 °C and concurrently, a tendency of the biopolymer component, with its increasing loading, to impede crystallinity development within the PCL phase (Figure 3). The higher Td value observed in AteCol film as compared to that of the AteCol porous matrix is attributed to a slightly low temperature dehydrothermal process occurring during preparation, which provides a low density of cross-links. Simultaneously, as determined from TG-DTG analysis (Figure 4), the Td large peaks from the DSC plots are accompanied by gradual mass decrement, which amounts from 7.5 to 5.8 or 13% for unmodified atelocollagen and the modified one (CH1P30−15, 2; CH10P5−15, 2), function of hydrophilic HA and hydrophobic PCL content, respectively. For unmodified atelocollagen, the major weight loss (63%) occurred at temperatures higher than 266 °C, while for modified materials this took place above 300 °C (66−75% weight loss, depending on composition), more quickly and easily after 300 and 330 °C, respectively. The higher denaturation transition temperature and increased thermal degradation values of modified materials as compared to AteCol alone demonstrate the cross-linking efficiency. To gain a better understanding of the behavior of the prepared matrices, DMA measurements were used to evidence the structural transition in the collagen based materials with temperature, the method being more sensitive to subtle structural changes with effect on mechanical viscoelastic properties. As shown in Figures 5 and 6, the variation of
Figure 6. Effect of AteCol modification on viscoelastic properties of the resulted materials. Dynamic storage modulus E′ variation for (a) AteCol, 2; (b) CH1P5−15, 2; (c) CH10P5, 2; and (d) CH10P5−15, 2.
217 °C for CH1P5−15, 220 °C for CH10P5 and 225 °C for CH10P5−15, respectively) correlated with the denaturation process and glass transition of the modified collagen in the developed materials, respectively. The inset detail in Figure 5 demonstrates the DMA ability as a sensitive characterization method to evidence subtle system changes, such as multiple denaturation transition in collagen based materials, usually difficult to be observed even by use of performant techniques.28 For AteCol, 2, the peak related to the denaturation process is bimodal, most probably due to a previous defibrillation followed by water elimination and breakage of the bonds that stabilize the secondary structure of collagen. It becomes larger and multimodal for the hybrid materials, shape associated with a stiffness increase effect and a more complex structure, the presence of hyaluronan inducing material and protein fibrils heterogeneity. The second peak in tan δ plot is corresponding to an abrupt decrement in E′ values (Figure 6), phenomenon attributed to the denaturation of dry collagen, that is, its transition into gelatin. It corresponds to the melting (softening) of the crystalline triple helix induced by the thermal disruption of hydrogen bonds, as also revealed by the DSC curves and in agreement with literature.26 The E′ values of the modified materials are lower than for AteCol alone, but the modulus drops to a very low value for higher temperature, for hybrid matrices, especially when they are submitted to UV treatment (Figure 6). At a temperature above 250 °C for unmodified AteCol and higher than 270 °C for modified hybrid samples, a very steep drop of the tan δ loss factor takes place, attributed to the glass-to-rubber transition of the denatured dry protein. The inclusion of DMSHA and PCL in formulation is also manifested by a shift of this peak in tan δ plot, associated with the process of decomposition toward higher temperatures, resulting in a higher thermostability (Figure 5). The formation of covalent bonds during the applied crosslinking procedures is also confirmed by tensile mechanical test results. The stress−strain curves in Figure 7 characterize the impact of the cross-linking procedure and incorporation of DMSHA and PCL on dense matrices mechanics. One can observe that the addition of DMSHA and PCL, as well as their percentage increase or submission to UV treatment, commonly yield strain decrease and Young’s modulus increase. This behavior is considered to accompany a decrease of the average inter-cross-linking distance (Dic). The cross-linked chains are
Figure 5. Loss tangent registration (DMA) for (a) AteCol, 2 and (b) CH10P5−15, 2.
elastic (E′) and viscous (E″ = Ei′ × tan δ) moduli, as well as of their ratio (loss factor − tan δ) versus temperature, is highly affected by cross-linking parameters, that is, by the inclusion of DMSHA and PCL in the formulation, as well as by UV irradiation duration. The enhancement in thermal stability is evidenced by the shifting toward higher values of the peaks tan δ plots, situated in the range of 20−80 °C (55 °C for AteCol, 62 °C for CH1P5−15, 2) and 200−230 °C (208 °C for AteCol, 2268
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address how much and how quickly such materials absorb the water from their surroundings, characteristics of interest in biomedical applications. For tissue engineering, the swelling properties of the scaffold significantly influence cell behavior: adhesion, growth, and differentiation.29 For wound dressing, an appropriate water absorbing capacity and moisture vapor transmission rate are between the main requirements, to achieve the proper moisture balance in the repairing wound, that is, to prevent both hydration and desiccation of the repairing tissue.30 In practice, hydrogels with high water uptake capacity and fast equilibrium swelling are preferred due to their higher permeability and biocompatibility. For the investigated materials, the water uptake increases at the initial stage, and then reaches equilibrium swelling in approximately 10 min for porous matrices and for dense films with a low DMSHA content, while for dense films with higher amounts of DMSHA the equilibrium is reached only after 20−25 min (Figure 8: CH10P5−15, 2 and CH10P30−15, 2). The ESR reflects the water uptake ability of the material, which is related to network structure, that is, to the hydrophilicity of the material (solvation of network chains) and to the morphological characteristics (presence and stability of a porous structure in water, related to the filling of the pores by the solvent).31 Overall, ESR slightly decreases with DMSHA and PCL amount, a higher cross-linking density implying less water retention. As expected, the porous specimens have a much higher swelling capacity as compared to the dense ones. However, as shown in Figure 9, for high DMSHA and PCL contents, the high cross-linking density gave rise to a nearly similar ESR value, but to a higher swelling rate comparative to the films with appropriate composition. More, for porous samples, the addition of small amounts of DMSHA yields an important increase in water uptake, as compared to AteCol alone (Figure 9). This behavior was explained by a positive contribution to polymer polarity coupled with the peculiarities of the porous structure, more important than the cross-links formation between the macromolecular compounds (collagen− GAG intermolecular bonding to generate composite), which should diminish the swelling ability. To obtain information on the swelling mechanism, the constants n were calculated from the slopes of the log/log plots designed according to eq 6 from the experimental data for the portion with a linear time dependence of the fractional water uptake for all geometries (Mt/M∞ lower than 0.6):
Figure 7. Deformation curves of unmodified AteCol (a) and of some prepared hybrid films: (b) CH1P5−15, 2; (c) CH10P5, 2; and (d) CH10P5−15, 2.
broken successively by extension and the breaking occurs well at shorter Dic (lower Mc), the localization of the stress occurring more easily. The upturn at high strain is indicative of strain-induced crystallization/enhanced molecular orientation, which causes the high modulus at high strain and high tensile strength. Hence, it should be noted that cross-linking by combination of AteCol with HA and PCL-DI may enhance the mechanical properties of the final matrices, and thus both physicochemical properties and processing features are adjustable via the control of formulation and preparation conditions. Data on swelling kinetics and water vapor sorption capacity of the dense and porous matrices are comparatively depicted in Figures 8 and 9 and Table 3. These are important parameters to
Figure 8. Effect of composition on swelling kinetics for dense films.
log M t /M∞ = log k + n log t
(6)
where Mt is the mass of water absorbed at time t and M∞ is the mass of water absorbed at equilibrium. All calculated values were lower than 0.5, indicating a preferential swelling by Fickian diffusion.32 The dynamic water vapor sorption capacity analysis gave results in agreement with the swelling behavior (Table 3). The results offer an insight into the structure−function relationship of the obtained materials. They can be used to predict the ease by which molecules can be transferred through such membranes, envisaging application in wound dressings. The morphologies of dried gel samples were observed by SEM to evaluate the effect of parameters that may control network microstructure: formulation, overall concentration in feed recipe, cross-linking degree, and preparation approach. Their variation can give rise to morphologies that may be references to modulate or improve the mechanical and swelling
Figure 9. Comparative swelling behavior in porous matrices.
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Table 3. Moisture Vapor Sorption Characteristics as Determined from Sorption−Desorption Isotherms and Dynamic Sorption Capacity Measurements GAB analysisa,c a
b
sample
sorption capacity (% d.b.)
dynamic sorption capacity, RH-90% (%)
AteCol, 2 CH1P30−15, 2 CH1P5−15, 2 CH10P5−15, 2 CH10P10−15, 1 CH1P2−30, 1
40.82 26.33 32.23 31.62 35.34 58.20
56.6 29.1 30.6 34.5 36.4 78.4
−3
b
sorption rate (×10 9.99 6.96 12.0 9.03 9.82 16.88
%/s)
2
AGAB (m /g)
monolayer (g/g)
436.485 355.046 384.225 393.299 422.531 393.862
0.12432 0.10113 0.10944 0.11202 0.12035 0.11218
Main parameters calculated from moisture-vapor sorption−desorption isotherms. bDeterminations performed at 25 °C until constant sample weight was reached for a relative humidity (RH) of 90% cSurface area and monolayer values calculated according to Guggenheim-Anderson-de Boer (GAB) model33−35 (IGAsorp soft). a
Figure 10. Typical SEM micrographs of porous and dense matrices. Samples: (a) AteCol, 1; (b) CH1, 1; (c) CH1P2−30, 1; (d) CH1ENP2−30, 1; (e) CH10P10−15, 1; (f) AteCol, 2; (g) CH1P5, 2; (h) CH10P5, 2; (i) CH1P30, 2; (j) CH10P30, 2.
decrease from 148 to 11 μm with increasing cross-linker percent in feed, while the thickness of pore wall increases, the hydrogel becoming more solid, especially with the amount of hyaluronate, when the walls present a cauliflower-like structure (Figure 10e). It is also obvious that the higher the polymer concentration in feed dispersion, the smaller the pore diameters and the thicker the pore walls. The specific architecture for high DMSHA content is most probably due to the aggregation of AteCol−DMSHA biocomposite microparticles (visible as spheres with 400−800 nm diameters), stabilized by PCL intra- and intercrosslinks. As observed earlier,11 high irradiation duration gives rise to an increase in pores size and heterogeneity, with pore wall disruption, as shown for sample CH1ENP2−30, 1, where the pores have diameters up to about 400 μm (Figure 10d). The observed high porosity of the lyophilized hydrogels and the interconnected pores ranging from 11 to about 200 μm are most striking, since these features are coincidentally preferred in their use as biomaterials. The possibility to control the pore size gradient by composition is beneficial for the use of the developed materials as scaffolds in tissue engineering, considering the high specificity of cells requirements. In contrast, with the increase in PCL and DMSHA content the uniform, smooth appearance of the film samples is changing to a heterogeneous, heteroporous one, with a more compact microstructure, with smaller and isolated pores as compared to
properties of the developed materials. As shown in Figure 10, the morphology of the samples essentially depends on the preparative approach. All lyophilized samples have a porous structure, while those formed by gradual drying under vacuum exhibit a continuous morphology without or with very few pores with irregular structure, no matter if they are cross-linked or not. The porosity for the dense films varies between 54 and 67%, while for the sponge matrices the porosity reaches values from 70 to 98%. The SEM microphotograph for un-crosslinked AteCol sponge sample clearly shows thin collagen fibrils and very few lattice-like lamellae, while the film sample has a smooth, uniform appearance (Figure 10a,f). For the porous specimens, the microstructure distinctly varies with formulation modification. It seems that, in the case of AteCol, the solution system was actually in a quasiequilibrium thermodynamic state, which could induce the collagen molecule to self-assembly longitudinally and laterally into long, relatively symmetrical fibrils (Figure 10a). By hyaluronan addition, the macromolecules accumulate into fine sheet-like structures, folding to generate pores with different sizes and forms (Figure 10b−d). Although the morphology is generally heteroporous, one can identify regular assemblies of interconnected pores with increased homogeneity for crosslinked samples (CH1P2−30, 1), especially when the crosslinker amount increases (CH10P10−15, 1), as compared to the un-cross-linked ones. The average pore diameter was found to 2270
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undestroyed even after 10 months for samples with PCL content higher than 10%. The difference in degradation rate depending on cross-linking degree is more evident during all process duration, the water access to heterogeneous bonds being determined by the combined effect of the hydrophobicity, crystallinity, chains flexibility (cross-linking degree), and bulk sample dimensions. With respect to the last factor, it seems that the porous samples are more accessible than the dense ones, with effect on the first stage of the degradation process. As previously reported,36 in cross-linked biomaterials, the degradation profile is controllable by tailoring sample composition and microstructure. Further results of this examination together with data from Fourier transformed infrared spectroscopy (FTIR-ATR) will be the subject of another publication focusing on degradation mechanism and main factors of influence. It must be mentioned that the hydrogels with a ternary composition are much more stable than those containing only AteCol and PCL,11 which, according to preliminary studies, are totally degraded during 3−24 h, even in PBS. The hemocompatibility assay showed that the formed 3D hybrid networks were non hemolytic, demonstrating their potential in serving as blood-contacting biomaterials. The values for hemolysis percentage (%) for all investigated samples were in the 0.002−0.02 range. Preliminary data on gentamicin sulfate upload and delivery point on an improved behavior as result of DMSHA and PCL inclusion in feed formulation.
the lyophilized samples. Generally, the pores are in the range of 5.5−10 μm for CH1Py samples (y = 5 or 30) and of about 12 μm, interconnected, for the CH10Py ones. In the last case, one may observe collagen fibrils and Col−GAG composite microparticles embedded in AteCol walls, densely connected or even covered by PCL. For a high percent of PCL in the hybrid films, as supposed, the excess of synthetic polyester is yielding phase separation and crystallization, in agreement with DSC data (Figure 10i,j). The design of new biomaterials envisaging possible use in biomedical area, especially of those related to tissue engineering, should not only be based on an adequate selection of the material but also on the capacity of previewing its possible behavior once implanted in terms of degradation mechanisms and feasibility of its performance in time. Once the tissue has taken over the synthetic support, the scaffold is expected to degrade in a controlled manner to nontoxic byproduct to match the new tissue growth rate. An often used solution is the inclusion of PCL in the composition, taking advantage of its long biodegradation, which allows tailoring of this characteristic for desirable biologic response. Durability during service life is also important for application as a component in wound dressings. Due to their composition, the obtained hydrogels are both hydrolytically (via ester group hydrolysis) and enzymatically degradable. To study their stability and better understand the mechanism, considering the complexity of the system, the degradation experiments were performed for comparison in two different media: alkaline (0.1 N NaOH) and phosphatebuffered solutions. The rate of degradation was higher for uncross-linked than for cross-linked materials. The durability was increased in both media with increasing PCL-DI and DMSHA amount in feed, behavior attributed to a higher cross-linking density (Figure 11). However, the initial degradation rate is
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CONCLUDING REMARKS To expand the range of available biomaterials, new hybrid natural/synthetic polymer based hydrogels comprising atelocollagen I, dimethylsilanediol hyaluronate and poly(ε-caprolactone), as sponges or dense films were prepared via simple and safe alternatives. The 3D products were characterized by FT-IR, WAXD, DSC, TGA, DMA, and SEM techniques supplemented by investigations on swelling behavior, water vapor sorption, degradation in alkaline, and PBS solutions and hemolytic tests. The obtained results proved that the used bifunctional poly(ε-caprolactone) is an efficient cross-linking agent even for the collagen−glycosaminoglycan system, and, consequently, it may be successfully applied in double/complex cross-linking of biopolymers, for the development of stable biodegradable, nontoxic, nonhemolytic materials. The introduction of DMSHA was shown to improve the swelling and water−vapor sorption capacity, while the addition of PCL in feed allows the control of durability and mechanical properties of the formed hydrogel, especially for the biocomposite films, without denaturation of collagen native structure, thus, combining the specific properties of natural and synthetic polymers incorporated therein. The high functionality and variety of the components makes possible the modification (extension) of the preparative protocol to develop new complex materials, that is, to promote the efficient inclusion of inorganic fillers, yielding polymer nanocomposites. Therefore, by optimizing the composition and preparative approach, materials of tailored physical, biological, and mechanical properties, as well as predictable degradation behavior, suitable for wound dressing (semiocclusive or biologic type) or periodontal membrane may be designed. Considering the exceptional tenability of the investigated hybrid materials, further studies focusing on the synthesis, characterization, and in vitro/in vivo testing of cryogels with a
Figure 11. Weight loss (%) of different scaffold compositions vs time at 37 °C in NaOH 0.1 N (a, b) and in PBS (c−e). Samples: (a) AteCol, 2; (b, c) CH1P5−15, 2; (d) CH10P10−15, 1; (e) CH10P30− 15, 2.
quite different, depending on solution characteristics and material composition. In alkaline media the degradation is highly accelerated, most probably beginning by scission of urethane, followed by ester bonds, giving rise to a rapid material fragmentation, followed by biopolymers fragments diffusion from the bulk material. Degradation rates are nearly similar. The most important factor of influence seems to be the PCL content. A different behavior was observed in the samples immersed in PBS when compared with those treated in NaOH. The PCL cross-links being more resistant, the collagen and MDSHA are less accessible and the network is disrupted at a slower rate, more than 30% of the network remaining 2271
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similar ternary composition, developed here for the first time, are in progress.
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(31) Rehakova, M.; Bakos, D.; Vizarova, K.; Soldan, M.; Jurickova, M. J. Biomed. Mater. Res. 1996, 30, 369−72. (32) Rittger, P. L.; Peppas, N. A. J. Controlled Release 1987, 5, 37−42. (33) Guggenheim, E. A. Application of Statistical Mechanics; Clarendon Press: Oxford, 1966; pp 186−206. (34) Anderson, R. B. J. Am. Chem. Soc. 1946, 68, 686−691. (35) de Boer, J. H. The Dynamical Character of Adsorption, 2nd ed.; Clarendon Press: Oxford, 1968; pp 200−219. (36) Hő glund, A.; Odelius, K.; Hakkarainen, M.; Albertsson, A.-C. Biomacromolecules 2007, 8, 2025−2032.
AUTHOR INFORMATION
Corresponding Author
*Phone: (+40) 232 217454. Fax: (+40) 232 211299. E-mail:
[email protected]. Notes
The authors declare no competing financial interest.
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