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Dynamic viscoelastic properties of collagen gels with high mechanical strength
Materials Science and Engineering C 33 (2013) 3230–3236
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Dynamic viscoelastic properties of collagen gels with high mechanical strength Hideki Mori, Kousuke Shimizu, Masayuki Hara ⁎ Department of Biological Science, Graduate School of Science, Osaka Prefecture University, 1-2 Gakuen-cho, Naka-ku, Sakai, Osaka 599-8570, Japan
a r t i c l e
i n f o
Article history: Received 6 July 2011 Received in revised form 8 March 2013 Accepted 29 March 2013 Available online 6 April 2013 Keywords: Collagen gel Storage modulus Loss modulus Young's modulus EDC Gamma-ray
a b s t r a c t We developed a new method for the preparation of mechanically strong collagen gels by combining successively basic gel formation, followed by 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) cross-linking and lyophilization. Gels cross-linked three times with this method showed stronger mechanical properties (G′: 3730 ± 2060 Pa, G″: 288 ± 35 Pa) than a conventional gel that was sequentially cross-linked with EDC once (G′: 226 ± 70 Pa, G″: 21 ± 4.4 Pa), but not as strong as the same gel with heating for 30 min at 80 °C (G′: 7010 ± 830 Pa, G″: 288 ± 35 Pa) reported in our previous paper. The conventional collagen gel was cross-linked with EDC once, heated once, and then subjected twice to a lyophilization–gel formation– cross-linking cycle to give three-cycled gel 2. This gel had the strongest mechanical properties (G′: 40,200 ± 18,000 Pa, G″: 3090 ± 1400 Pa, Young's modulus: 0.197 ± 0.069 MPa) of the gels tested. These promising results suggest possible applications of the gels as scaffolds in tissue engineering research. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Collagen is an extracellular matrix (ECM) protein. It is biocompatible and is found throughout the connective tissues of the bodies of animals. It has been used to fabricate gels and sponges for medical use as scaffolds in tissue engineering and for medical devices [1]. Collagen gels have also been used as matrices for drug delivery systems (DDSs) [2]. Collagen gels can be cross-linked by various methods to mechanically reinforce the gel, and also to control the rate of bioabsorption (degradation in vivo). However, the combination of mechanical strength and elasticity of the cross-linked collagen gels has so far not been adequate for applications in medical devices, especially in the clinical fields of orthopedics, cardiovascular surgery, and neurosurgery. It is clear that collagen gels of higher strength are needed [1,3–5]. Various methods for cross-linking, based on different principles, have been developed [1]. They can be broadly classified as i) natural cross-linking between hydroxyl lysine residues [1], ii) dehydrothermal cross-linking under vacuum at high temperatures, which forms intermolecular cross-links through condensation reactions, either esterification or amide formation [2,6], iii) treatment with glutaraldehyde (GTA) to effect crosslinking between two amino residues (\NH2) [1,2], iv) treatment with a carbodiimide (e.g., EDC) to crosslink between an amino residue (\NH2) and a carboxyl residue (\COOH) [1,7], v)
Abbreviations: EDC, 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide. ⁎ Corresponding author. Tel./fax: +81 72 254 9842. E-mail address: [email protected] (M. Hara). 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.03.047
other synthetic cross-linkers [1,2,8], vi) enzymatic reactions using transglutaminase [9], vii) other natural compounds as cross-linkers [10–12], viii) metal ions such as chromium(III), aluminum(III), zirconium(IV), iron(II) [13,14] and ix) radiation-based cross-linking through radical reactions (e.g., UV light, dye-mediated photooxidation, gamma rays, and electron-beam irradiation [1,5]). There is a continuing need for new cross-linking methods for collagen gels to reinforce the mechanical properties of the gel and avoid any toxicity due to residual reagents. From the various methods listed above, chemical cross-linking with GTA and with EDC has probably been the most widely used in tissue engineering and DDSs, although other cross-linkers of lower toxicity have emerged recently [1,2]. We therefore chose EDC as the cross-linker for the present study. We measured the dynamic viscoelasticities of collagen gels prepared and modified by four different methods: i) collagen gels crosslinked by EDC after preparation, ii) collagen gels cross-linked simultaneously with their preparation, iii) collagen gels irradiated with gamma rays after preparation, and iv) collagen gels formed directly from an acidic collagen solution by gamma cross-linking, as described previously [15]. We measured their mechanical properties using tensile testing and also the viscoelastic properties. From our results, we conclude that collagen gels with highly cross-linked thick collagen fibrils after heating showed a higher value of storage modulus (G′) and loss modulus (G″), compared with the previous report [15]. Using the experience obtained from our previous work, we succeeded in preparing collagen gels with extremely high mechanical strength by repeating the cycle of gel formation (fibril formation), cross-linking with EDC, and freeze-drying in which the heating was done only once after cross-linking in the first cycle. We also discuss the possible future uses of these materials.
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2. Materials and methods Materials and methods, including reagents, SEM observations, measurements of dynamic viscoelasticity, and tensile tests, were essentially as described in our previous paper [13]. We briefly describe the materials and methods. Porcine type I collagen solution (0.6% (w/v), pH 3.0, collagen BM, Nitta Gelatin, Osaka, Japan), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) (water-soluble carbodiimide hydrochloride 1030, Wako, Osaka, Japan), and other reagents were of a specific grade or analytical grade, as noted. A rheometer with an oscillatory rotating disk (Physica MCR301, Anton Paar GmbH, Graz, Austria) was used for dynamic rheological measurements. A tensile tester (Autograph AGS-50ND, Shimadzu, Kyoto, Japan) was used for the tensile tests. A freeze-dryer (FD-1, Eyela, Tokyo, Japan) and a deep freezer (Ultra Low, Sanyo, Tokyo, Japan) were used to lyophilize the gel. An electronic digital external micrometer (MCD130-25, Niigata Seiki, Sanjo, Japan) was used to measure the thickness of the collagen gel disks. An ion coater (IB-3, Eiko Engineering, Hitachinaka, Japan) and a scanning electron microscope (SEM; SM-100, Topcon, Tokyo, Japan) were used to characterize the gels. The preparation of the gels using cycled gel-formation is described below. Collagen gels were prepared by pH neutralization. Collagen (4 mL, 0.6% w/v), 0.8 mL of 10-fold concentrated phosphate-buffered saline (PBS), and 3.2 mL distilled water were mixed in a 6-well plate and incubated overnight at 37 °C to form the collagen gels. These were then incubated overnight in 10 mL PBS containing 50–125 mM EDC. The gels were washed well with distilled water to remove the salts. Each sample was placed on a plastic tray (35 mm in diameter, 20 mm in depth) and then wrapped with a thin plastic film, placed in a Styrofoam box (10 × 15 × 20 cm in length × width × depth) and frozen in a deep freezer at −85 °C overnight. The frozen sample was then lyophilized in a freeze-dryer to form the chemically cross-linked collagen sponge (microporous material). The sponge was immersed in the ice-cold collagen solution just after neutralization of the solution by addition of PBS. The solution promptly permeated into
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the sponge at 4 °C. The sample was then warmed and incubated at 37 °C to form the collagen gels inside the micropores in the sponge. The sample was incubated overnight in PBS containing 50 mM EDC for cross-linking. By repeating the cycle of collagen gel-formation followed by EDC cross-linking and lyophilization three times, we prepared what we refer to as the “3-cycle collagen gel 1” sample. A second mixture, consisting of 4 mL 0.6% collagen, 3.2 mL of distilled water, 0.8 mL of × 10 PBS(−), was prepared in a 6-well plate. After neutralization, the collagen gel was cross-linked with EDC and then heated for 30 min at 80 °C. The volume of the sample decreased significantly during the heating process. This gel is referred to as “1-cycle collagen gel”. The sample was lyophilized to form a sponge and was used to prepare a collagen gel followed by cross-linking with EDC twice. This gel is referred to as the “3-cycle collagen gel 2” in this paper. If the EDC-treatment was conducted only once, the gel was termed “2-cycle collagen gel 2”. Conditions for preparation of the various gels are shown with abbreviations in Tables 1 and 2 and in our previous paper [15]. All samples were prepared as described above and used for the rheological measurements in a wet state without additional lyophilization, and therefore are described as “gels”, not as “sponges”, in this paper and our previous paper [15]. Collagen gel disks (diameter, φ = 8–10 mm, thickness = 1–2 mm) were prepared and used for dynamic viscoelasticity measurements using a rheometer with an oscillatory rotating disk to obtain values for the storage modulus (G′) and the loss modulus (G″). Measurements were carried out with a deflection angle of 0.167–0.501 m rad and a range of angular frequency ω from 0.1 to 100 rad/s at ambient temperature (approximately 25 °C). Four samples (n = 4) were measured for each set of conditions, and the mean and standard deviation were calculated. The value of tangent δ (tan δ) was calculated as the ratio G ″/G′. This parameter reflects the thermal energy loss. The G′ value at 1 rad/s was assumed to be the G value. The number of network points per cubic meter (ν) was calculated using the formula: G ¼ νkT
Table 1 Summary of G′ and G″ values from the measurements of dynamic viscoelasticity. Abbreviations are Gel = gel-formation by neutralization of pH, H = heat-treatment for 30 min at 80 °C, 50E = the gel was treated with 50 mM EDC, and 125E = the gel was treated with 125 mM EDC. *Data recorded using conditions from Table 1 of the previous paper [13]. Curves, with standard deviations, are shown in Supplementary data 1–3. These values were slightly dependent on the range of rad/s (0.1–10) and frequency. Thus, values at the frequency of 1 rad/s are shown for comparison purposes in this table. Values of G′ and G″ are shown as means ± standard deviation (n = 4). If we compare pairs of unheated and heated gels, as analyzed by Student's t-test, no significant difference (p b 0.05) was found between the heated and unheated samples of any gel except between Gel/125E and Gel/125E + H (p b 0.01, as shown by #a). If we compare the heated sequential gel (Gel/125E + H) with both 2-cycled gel 2 and 3-cycled gel 2, values of viscoelastic parameters (G′ and G″) increased significantly with each cycle of gel formation. The superscript symbols with letters (#; p b 0.01, +; 0.01 ≤ p b 0.05) indicate significant differences between pairs of unheated and heated gels. Type of gel Abbreviation i) Uncross-linked collagen gel Gel*
G′ (Pa)
tan δ
G″ (Pa)
44.1 ± 31.3
ν (×1023, number/m3)
11.7 ± 8.8
0.265
0.107
ii) Sequentially EDC-cross-linked collagen gel (sequential gel) Gel/125E* 226 ± 70#a Gel/125E + H* 7010 ± 830#a,+c,+e,+g,#i
21 ± 4.4#b 288 ± 35#b,+d,#f,#h,#j
0.091 0.041
0.549 17.0
iii) 3-cycled collagen gel 1 (Gel/125E) × 3 times (Gel/125E) × 3 times + H
288 ± 35 168 ± 27
0.054 0.066
9.01 6.16
3730 ± 2060 2550 ± 517
iv) 2-cycled collagen gel 2 (Gel/50E) + H + (Gel/50E) (Gel/50E) + H + (Gel/50E) + H (Gel/125E) + H + (Gel/125E) (Gel/125E) + H + (Gel/125E) + H
25,400 16,200 20,400 26,100
± ± ± ±
13,700 2680 9780+c 11,700+e
842 657 944 1170
± ± ± ±
497 199 506+d 399#f
0.033 0.041 0.046 0.045
61.4 39.1 49.5 63.5
v) 3-cycled collagen gel 2 (Gel/50E) + H + (Gel/50E) × 2 times (Gel/50E) + H + (Gel/50E) × 2 times + H (Gel/125E) + H + (Gel/125E) × 2 times (Gel/125E) + H + (Gel/125E) × 2 times + H
18,000 16,800 40,200 38,500
± ± ± ±
6070 2800 18,000+g 6740#i
1300 765 3090 1920
± ± ± ±
416 94 1400#h 329#j
0.072 0.046 0.077 0.050
43.8 40.7 97.6 93.6
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Table 2 Summary of result from the tensile tests. Maximal values of the tensile stress (σ), tensile strain (ε) and corresponding Young's modulus (E) are shown as mean ± standard deviation (n = 7). *Data recorded using parameters from Table 1 of the previous paper [13]. Abbreviations are the same as in Table 1. There was no significant difference (p b 0.05) between the results from the sample with additional heating at the end (+H) and the corresponding unheated sample for each type of 3-cycled gel, as analyzed by Student's t-test. Comparing sequential gels, results were significantly different before and after the final heating step (+H) except for the values of ε for Gel/50E and Gel/50E + H, as described in Table 3 of the previous paper [13]. If we compare the results from the 3-cycled collagen gel 2 with those of sequential gels, the difference was significant. The superscript symbols with letters (#; p b 0.01, +; 0.01 ≤ p b 0.05) show that there were significant differences between the pair of unheated and heated gels. Type of gel Abbreviation
Tensile stress σ (MPa)
Tensile strain ε
Young's modulus E (MPa)
ii) Sequentially EDC-cross-linked gel (sequential gel) Gel/50E* Gel/50E + H*
0.010 ± 0.008#n 0.066 ± 0.019+a,#d,#n
4.37 ± 0.91 4.19 ± 1.28#b,#e
Gel/125E* Gel/125E + H*
0.004 ± 0.001#q 0.053 ± 0.011#h,#k,#q
3.34 ± 1.33 5.29 ± 1.50#i,#l
0.0047 0.013 0.097 0.0034 0.012 0.095
± ± ± ± ± ±
0.004#o,#p 0.0044#f,#o 0.035#c,#g,#p 0.0014#r,#s 0.005#j,#m,#r 0.034#s
v) 3-cycled collagen gel 2 (Gel/50E) + H + (Gel/50E) × 2 times (Gel/50E) + H + (Gel/50E) × 2 times + H (Gel/125E) + H + (Gel/125E) × 2 times (Gel/125E) + H + (Gel/125E) × 2 times + H
0.112 0.137 0.129 0.164
0.135 0.179 0.197 0.219
± ± ± ±
0.068#c 0.031#f,#g 0.069#j 0.051#m
± ± ± ±
0.050+a 0.035#d 0.040#h 0.063#k
where T is the absolute temperature (298 K) and k is the Boltzmann constant (1.38 × 10−23 J/K). Collagen gel disks, prepared as described, were pressed between two plastic sheets to a final size of φ = 8–10 mm, thickness = 1– 2 mm, and then subjected to tensile tests at ambient temperature (approximately 25 °C) using a tensile tester with an extension speed of 1 mm/min. Two pieces of sandpaper were placed on the surfaces of the sample folder to prevent slippage of the gels. Seven (n = 7) samples in each condition were prepared for the tensile test. The section area (A = thickness × diameter (mm 2)) of each sample was calculated from values measured using a digital micrometer. We did the tensile tests to obtain a stress (N)–strain (mm) curve. The original length of the sample in the tensile test (representing the distance between the upper and lower holders) was fixed at 1 mm. For a change in l of Δl, the tensile strain (ε) = Δl/l. The value of ε is thus the ratio between the original length and the change in length, and is therefore dimensionless. The tensile stress (σ) (MPa) = stress (N)/A (mm 2). Young's modulus (E) values were calculated from the maximal values of these parameters and the cross-sectional area of the sample gel, to give E = σ/ε. Various types of collagen gel, prepared in a 24-well plate, were washed with PBS and fixed for 3 h in PBS containing 2.5% (w/v) glutaraldehyde at 4 °C. The fixed samples were washed for 1 h in PBS, and then incubated successively for 20 min each in 30%, 50%, 70%, 80%, and 90% (v/v) ethanol at 4 °C. The samples were incubated for 20 min twice at ambient temperature to dehydrate them. The samples were then incubated for 20 min twice in a mixture of t-butanol and ethanol (1:1) to exchange the solvent, frozen in a freezer, and then dried under vacuum in a freeze-dryer. The dried samples were coated with platinum using an ion-coater and then observed using a scanning electron microscope (SEM). 3. Results We consider that both the thickness of the collagen fibrils and their 3D-density are important parameters for defining the viscoelastic properties of the collagen gels. Usually, these two parameters depend on the concentration of collagen used for the preparation of the gels. However, the acidic solution of 0.6% (w/v) collagen is already very viscous. It seems difficult to increase the concentration of the solution. We therefore invented a new method using repeated gel-formation inside the micropores of a collagen sponge. The collagen gels were prepared, cross-linked with EDC, washed and lyophilized, and then immersed in the collagen solution just before gel
1.26 1.14 1.06 1.07
± ± ± ±
0.32#b 0.28#e 0.21#i 0.23#l
formation to increase the 3D-density of the collagen fibrils in the gel. The number of cycles used is noted in Table 1, and described in the Materials and methods section, above. The shrinkage of the gels was quite distinct in the first heating cycle but less so in the second and third heating cycles. The three samples were used for SEM observations and rheological measurements of the dynamic viscoelasticity. The 2-cycled gel was white and non-transparent (data not shown). Fig. 1 shows SEM images of the 2-cycled gel prepared under various conditions. Collagen fibrils are observed in the gel treated with 50 mM EDC (A). These were twisted and fused to each other, forming thick bent fibrils with the inter-fibril space, reduced by additional heating (B). Thicker fused collagen fibrils were observed in the 2-cycled gels treated with 125 mM EDC before (C) and after the additional heating (D). The 3-cycled gel was also white and non-transparent, like the 2-cycled gel (data not shown). No significant change in size or appearance was observed after additional heating. Fig. 2 shows SEM images of the 3-cycled gel. The sample treated with 50 mM EDC showed a rough surface with fuzzy fibrils on the surface, as shown before (A) and after (B) heating. For the sample treated with 125 mM EDC, the fibrils were more densely packed and individual fibrils were difficult to observe, as in (C). The sample after heating showed a relatively smooth surface, as in (D). The interfibril spaces in (C) and (D) are believed to be small because of the densely packed collagen fibrils. The G′ and G″ values obtained from the measurements of dynamic viscoelasticity are summarized in Table 1. As mentioned previously, the gel with the highest values of G′ and G″ was the 3-cycled gel ((Gel/125E) + H + (Gel/125E) × 2 times). The second highest was that with additional heating ((Gel/125E) + H + (Gel/125E) × 2 times + H). The third highest was the 2-cycled gel ((Gel/125E) + H + (Gel/125E) + H). All of the 3-cycled gel 2 samples had higher values of G′ and G″ (G′ > 16,000 Pa, G″ > 700 Pa) than those of the 3-cycled gel 1 (G′ b 4000 Pa, G″ b 300 Pa) without heating after their first cross-linking ((Gel/125E) × 3 times, and (Gel/125E) × 3 times + H). The gel cross-linked once with EDC without heating (Gel/125E) showed higher values (G′; 226 Pa, G″; 21 Pa) than the uncrosslinked gel (Gel) (G′; 44.1 Pa, G″; 11.7 Pa). However those values were far lower than those (G′; 7010 Pa, G″; 288 Pa) of the sample with heating (Gel/125E + H). As described in our previous paper, the first heating made the gel shrink, causing the collagen fibrils to bend, twist, and fuse to each
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A
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B
2.0 µm
C
2.0 µm
D
2.0 µm
2.0 µm
Fig. 1. SEM images of the 2-cycled gels, (Gel/50E) + H + (Gel/50E) (A), (Gel/50E) + H + (Gel/50E) + H (B), (Gel/125E) + H + (Gel/125E) (C), and (Gel/125E) + H + (Gel/ 125E) + H (D). The images show the surface of the gels.
other [15]. We conclude that these changes brought about by heating contributed to the increase of G′ and G″ values, in addition to the cross-linking. We also need to understand the effect of the final heating on the cycled and non-cycled gels. The gel, cross-linked once with 125 mM EDC, showed a distinct increase in G′ value, from 226 to 7010, by
heating. However, a 2-cycled gel cross-linked with 125 mM EDC showed only a slight increase in the G′ value, from 20,400 to 26,100, after the final heating while the 3-cycled gel cross-linked with 125 mM EDC showed a decrease in G′ value, from 40,200 to 38,500. Thus, it appears that the strengthening effect brought about by heating after cross-linking is a distinct process after the first
Fig. 2. SEM images of the 3-cycled gels, (Gel/50E) + H + (Gel/50E) × 2 times (A), (Gel/50E) + H + (Gel/50E) × 2 times + H (B), (Gel/125E) + H + (Gel/125E) × 2 times (C), and (Gel/125E) + H + (Gel/125E) × 2 times + H (D). The images show the surface of the gels.
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in the same direction, parallel to the extension, and therefore show low tensile stress. In the moderately increasing phase, bundles of the thick collagen fibrils are mechanically strained, showing higher tensile stress. We expect that the shapes of the curves for the 3-cycled gel 2 shown in Fig. 3 of this paper can be described similarly. A network of thick twisted collagen fibrils (diagram B in Fig. 7 of Ref. [15]) and that of fibrils of intermediate thickness (diagram C in Fig. 7 of Ref. [15]) become interpenetrated and entangled with each other. Both types of fibrils become highly crosslinked with EDC. Thus, bundles of thick collagen fibrils are mechanically strained, showing higher tensile stress than the moderately-increasing phase. The rapidly increasing phase was diminished because almost all the heat-denatured collagen (gelatin) was already crosslinked with EDC. Data obtained in the tensile test are summarized in Table 2. The values of Young's modulus were somewhat variable, falling within the range of the 1.3-fold change in the sample treated with both 50 mM and 125 mM EDC with and without heating. There was no significant difference (p b 0.05) between the heated sample and unheated one for each set of conditions, as analyzed by Student's t-test. From the data in Table 2, we conclude that all four types of 3-cycled gel 2 were mechanically strong. The tensile strain was in the range from 1.26 to 1.06, so we also conclude that the 3-cycled gel 2 samples have low tensile strain.
cross-linking, but becomes less apparent after further cross-linking cycles. With 50 mM EDC, even the 2-cycled gel showed a decrease in the G′ value with heating, from 25,400 to 16,200, suggesting the possibility that uncross-linked collagen had been partially dissolved. There was no significant difference (p b 0.05) between the heated sample and unheated one for each set of conditions in Table 1, as analyzed by Student's t-test, except the comparison between Gel/125E and Gel/125E + H. Thus, we conclude that the final heating process did not greatly alter the rheological properties. Examples of the tensile test data (tensile strain–tensile stress curve) for the 3-cycled gel 2 are shown in Fig. 3. The shapes of the curves for gels prepared using four different conditions (Fig. 3(A)–(D)) were similar to each other. The initial increase in the tensile stress was almost linear and then the stress gradually reached a plateau. The maximal tensile stress at a final breaking point was variable, and in the range 0.05–0.25. The strain at the maximal tensile stress was typically small (ε b 1.5) in the sample treated with 125 mM EDC, suggesting significant stiffness of the sample. We measured the tensile test of the sequentially EDC-cross-linked collagen gel, as shown in our previous paper [15]. The shapes of the curves for the unheated sequentially EDC cross-linked collagen gel consist of three phases – a lag phase, a moderately increasing phase, and a plateau – while another rapidly-increasing phase was present for the heated gel, as shown in Fig. 3 of our previous paper [15]. By comparison with these data, the curves for the present 3-cycled gels clearly have neither a lag phase nor a rapidly-increasing phase. Why the shapes of the curves are so different between the cycled gel 2 in this paper and the sequential gel in the previous paper [15] is not easy to answer but seems to relate to the mechanism of deformation of the gel in the tensile test. We believe that the rapidly-increasing phase in the heated sample of sequential gels is derived from the heat-denatured collagen (gelatin), as described in the previous paper [15]. We now consider the shapes of the curves in the sequential gel using the previous arguments [15]. The initial lag phase in the unheated sequential gels corresponds to the process in which the cross-linked collagen fibrils are mechanically forced to be oriented
4. Discussion We succeeded in preparing a mechanically durable collagen gel using successive cycles of gel formation by neutralization, cross-linking with EDC, lyophilization, and heating. We determined that the 3-cycled collagen gel 2, processed with 125 mM EDC but without heating, was mechanically very strong (storage modulus G′; 40,200 Pa, loss modulus G″; 3090 Pa, Young's modulus E; 0.197 MPa, tensile stress σ; 0.129 MPa, tensile strain ε; 1.06) and shows promise as a scaffold for tissue engineering research. This 3-cycled gel 2 was mechanically the strongest collagen gel we have prepared. However,
A
B 0.3
Tensile stress MPa
Tensile stress MPa
0.3
0.25 0.2
0.15 0.1
0.05
0.25 0.2 0.15 0.1 0.05 0
0
0
0.5
1
1.5
2
0
Tensile strain
1
1.5
2
Tensile strain
C
D 0.3
0.3
Tensile stress MPa
Tensile stress MPa
0.5
0.25 0.2 0.15 0.1 0.05 0
0
0.5
1
Tensile strain
1.5
2
0.25 0.2 0.15 0.1 0.05 0 0
0.5
1
1.5
2
Tensile strain
Fig. 3. Examples of the tensile stress–tensile strain curves obtained in the tensile test of (Gel/50E) + H + (Gel/50E) × 2 times (A), (Gel/50E) + H + (Gel/50E) × 2 times + H (B), (Gel/125E) + H + (Gel/125E) × 2 times (C), and (Gel/125E) + H + (Gel/125E) × 2 times + H (D). Different line types show different examples of results using the same measurement conditions.
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the 3-cycled gel 2 was stiff and not tensible because the tensile strain was only 1.06. We considered why the cycled gels were mechanically much stronger than conventional collagen gels crosslinked once with EDC but show only moderate increases in the values of their Young's modulus. However, we currently have no clear answer to this question. It is possible that a tiny crack formed in the sample, causing it to tear during the tensile test, and so reduce the maximal tensile stress, finally resulting in a moderate value of Young's modulus. As also discussed in our earlier paper [15], a mechanically strong scaffold is sometimes necessary, in the tissue engineering of bone, cartilage, tendon, ligament, muscles, and other tissues. In particular, if we culture chondrocytes to repair injured cartilage, mechanically strong collagen gels or collagen sponges are definitely necessary [16]. The mechanical properties of tissues are determined by the density and orientation of the collagen fibrils, the extent of covalent cross-linking between the collagen molecules, and the contribution of other extracellular matrix (ECM) components, such as elastin, glycosaminoglycans, and proteoglycans. Collagen fibrils are in relatively random orientations in the dermis and cartilage, but are highly oriented, unidirectionally, in tendons, ligaments, and trabecular bone. Cross-linked networks of collagen and elastin are also contained in the laminated layers of the walls of blood vessels, providing their tensibility and elasticity. We prepared the 3-cycled collagen gel 2 with good mechanical strength, although the orientation of the collagen fibrils was not controlled in this research. We consider that a scaffold for cartilage is a candidate use of this gel, in tissue engineering [16]. However, the pore size appears to be small and will not permit cells to proliferate into the scaffold material. This can be advantageous or disadvantageous, depending on the application. This problem should be addressed by additional steps of lyophilization or some other method in the future. We used 0.6% (w/v) porcine atelocollagen that is commercially available. We did not try to concentrate the collagen solution further in this study. If the viscosity of the solution is too high, it is difficult to mix the solution with PBS for neutralization. It has been reported that inter-molecular interactions are present in very concentrated solutions of collagens [17]; liquid crystal-like nematic phases and precholesteric phases appear in 2% (w/v) and 4% (w/v) solutions, respectively. These intermolecular interactions may contribute to the mechanical properties of the collagen gel by affecting the thickness and orientation of the collagen fibrils. These issues remain to be further characterized. In the preparation of the gel with the highest G′ value ((Gel/ 125E) + H + (Gel/125E) × 2 times), the collagen gel was made into a porous collagen sponge by the first freeze-drying, with crystals of ice acting as porogens. Then, the second fibril-formation occurred in the pores (inner spaces) of the sponge thus forming the hierarchical structure of the collagen fibrils. The first fibrils to be formed were partially denatured into thick twisted fibrils or were fused together by the heating, while the second and third fibrils were not. The denatured thick collagen fibrils and undenatured fibrils show different mechanical properties, as described in our previous paper [15]. We assume that we made two collagen gels, with differing mechanical properties, even though we used the same raw material, collagen. It has also been reported that the pore size of the collagen sponge is an important factor governing cell migration into the sponge. If the pore size is too large, the surface area is small. If the pore size is too small, the cells find it difficult to migrate into the sponge. O'Brien and colleagues tested in detail the effects of various conditions (e.g., collagen concentration, freezing rate, temperature) on the preparation of the sponge by freeze-drying [18–20] and concluded that a mean pore size of 120 μm was the most suitable for the migration of cells into a collagen–glycosaminoglycan sponge [21]. We have
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not yet measured the pore size or tested the migration of cells into the cycled gels prepared in this study but this will be a subject for future investigations. The mechanical properties – especially the tensile strength – of the collagen gel fabricated also depend on the orientation of the collagen fibrils. It has also been reported that a collagen gel block, once prepared, can be compressed mechanically to a thin film, and then sequentially cross-linked with EDC to form an EDC-crosslinked collagen gel sheet [22–24]. The orientation of the collagen fibrils along the z-axis was not controlled, but neither was the orientation along the x- or y-axis. This method can effectively increase the three-dimensional density of the collagen fibrils. Sheets or films consisting of collagen gel are useful as structural biomaterials. They can be cut with scissors, laminated to form multi-layers [25], or rolled up to form concentric lamellae, like construction paper [22–24,26]. Zeugolis et al. prepared extruded collagen fibers in which collagen fibrils were oriented [27,28], and then treated them by various crosslinking methods, as listed in Table 2 of Ref. [28]. Tanaka et al. reported the preparation of collagen gel sheets in which the collagen fibrils were highly oriented, using an oriented flow casting method [26]. We have not yet tried to control the orientation of collagen fibrils in the present work but this will also be an interesting theme for future investigations. 5. Conclusions We have described the preparation of mechanically durable collagen gels (G′; 40,200 ± 18,000 Pa, G″; 3090 ± 14,000 Pa) by repeating the gel-formation/EDC cross-linking/lyophilization cycle. The gels prepared by various methods including cross-linking and heating, need to be further characterized in terms of the technical issues of cell adhesion properties, biocompatibility, and bioabsorptivity (non-toxic biodegradability in vivo) before their possible use in medical devices. Acknowledgments We thank Dr. Shiho Suzuki, and Shinichi Kitamura in Osaka Prefecture University for their support in the measurements of dynamic viscoelasticity, and to Ms. Hiroko Machida of AIST-Kansai for SEM observations. Appendix A. Supplementary data Supplementary data to this article can be found online at http:// dx.doi.org/10.1016/j.msec.2013.03.047. References [1] S.F. Badylak, in: A. Atala, R.P. Lanza (Eds.), Methods of Tissue Engineering, Academic Press, London, 2002, pp. 503–514. [2] W. Friess, Eur. J. Pharm. Biopharm. 45 (1998) 113. [3] I.V. Yannas, J. Macromol. Sci., Rev. Macromol. Chem. C7 (1972) 49B. [4] I.V. Yannas, Clin. Mater. 9 (1992) 179. [5] M. Hara, J. Oral Tissue Eng. 3 (2006) 118. [6] M.G. Haugh, M.J. Jaasma, F.J. O'Brien, J. Biomed. Mater. Res. 89A (2009) 363. [7] S. Yunoki, K. Mori, T. Suzuki, N. Nagai, M. Munekata, J. Mater. Sci. Mater. Med. 18 (2007) 1369. [8] E.C. Collin, S. Grad, D.I. Zeugolis, C.S. Vinatier, J.R. Clouet, J.J. Guicheux, P. Weiss, M. Alini, A.S. Pandit, Biomaterials 32 (2011) 2862. [9] J.M. Orban, L.B. Wilson, J.A. Kofroth, M.S. El-Kurdi, T.M. Maul, D.A. Vorp, J. Biomed. Mater. Res. 68A (2004) 756. [10] D.I. Zeugolis, R.G. Paul, G. Attenburrow, Mater. Sci. Eng. C30 (2010) 190. [11] F.-L. Mi, C.-T. Huang, Y.-L. Chiu, M.-C. Chen, H.-F. Liang, H.-W. Sung, J. Biomed. Mater. Res. 83A (2007) 667. [12] H.-W. Sung, W.-H. Chang, C.-Y. Ma, M.-H. Lee, J. Biomed. Mater. Res. 64A (2003) 427. [13] A.D. Covington, Chem. Soc. Rev. (1997) 111. [14] B. Wu, C. Mu, G. Zhang, W. Lin, Langmuir 25 (2009) 11905. [15] H. Mori, K. Shimizu, M. Hara, Mater. Sci. Eng. C32 (2012) 2007.
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[16] J.M. MacPherson, R. Tubo, in: R.P. Lanza, R. Langer, J. Vacanti (Eds.), Principles of Tissue Engineering, 2nd edition, Academic Press, London, 2000, pp. 697–709. [17] L. Besseau, B. Coulomb, C. Lebreton-Decoster, M.-M. Giraud-Guille, Biomaterials 23 (2002) 27. [18] F.J. O'Brien, B.A. Harley, I.V. Yannas, L.J. Gibson, Biomaterials 26 (2005) 433. [19] C.M. Tierney, M.G. Haugh, J. Liedl, F. Mulcahyy, B. Hayes, F.J. O'Brien, J. Mech. Behav. Biomed. Mater. 2 (2009) 202. [20] F.J. O'Brien, B.A. Harleyb, I.V. Yannas, L. Gibson, Biomaterials 25 (2004) 1077. [21] C.M. Murphy, M.G. Haugh, F.J. O'Brien, Biomaterials 31 (2010) 461. [22] R.A. Brown, M. Wiseman, C.-B. Chuo, U. Cheema, S.N. Nazhat, Adv. Funct. Mater. 15 (2005) 1762.
[23] M. Bitar, V. Salih, R.A. Brown, S.N. Nazhat, J. Mater. Sci. Mater. Med. 18 (2007) 237. [24] E.A.A. Neel, U.C. Jonathan, C. Knowles, R.A. Brown, S.N. Nazhat, Soft Matter 2 (2006) 986. [25] M. Hara, A. Yamaki, J. Miyake, Mater. Sci. Eng. C17 (2001) 107. [26] Y. Tanaka, K. Baba, T.J. Duncan, A. Kubota, T. Asahi, A.J. Quantock, M. Yamato, T. Okano, K. Nishida, Biomaterials 32 (2011) 3358. [27] D.I. Zeugolis, R.G. Paul, G. Attenburrow, J. Biomed. Mater. Res. 86A (2008) 892. [28] D.I. Zeugolis, G.R. Paul, G. Attenburrow, J. Biomed. Mater. Res. 89A (2009) 895.
Contents lists available at SciVerse ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Dynamic viscoelastic properties of collagen gels with high mechanical strength Hideki Mori, Kousuke Shimizu, Masayuki Hara ⁎ Department of Biological Science, Graduate School of Science, Osaka Prefecture University, 1-2 Gakuen-cho, Naka-ku, Sakai, Osaka 599-8570, Japan
a r t i c l e
i n f o
Article history: Received 6 July 2011 Received in revised form 8 March 2013 Accepted 29 March 2013 Available online 6 April 2013 Keywords: Collagen gel Storage modulus Loss modulus Young's modulus EDC Gamma-ray
a b s t r a c t We developed a new method for the preparation of mechanically strong collagen gels by combining successively basic gel formation, followed by 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) cross-linking and lyophilization. Gels cross-linked three times with this method showed stronger mechanical properties (G′: 3730 ± 2060 Pa, G″: 288 ± 35 Pa) than a conventional gel that was sequentially cross-linked with EDC once (G′: 226 ± 70 Pa, G″: 21 ± 4.4 Pa), but not as strong as the same gel with heating for 30 min at 80 °C (G′: 7010 ± 830 Pa, G″: 288 ± 35 Pa) reported in our previous paper. The conventional collagen gel was cross-linked with EDC once, heated once, and then subjected twice to a lyophilization–gel formation– cross-linking cycle to give three-cycled gel 2. This gel had the strongest mechanical properties (G′: 40,200 ± 18,000 Pa, G″: 3090 ± 1400 Pa, Young's modulus: 0.197 ± 0.069 MPa) of the gels tested. These promising results suggest possible applications of the gels as scaffolds in tissue engineering research. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Collagen is an extracellular matrix (ECM) protein. It is biocompatible and is found throughout the connective tissues of the bodies of animals. It has been used to fabricate gels and sponges for medical use as scaffolds in tissue engineering and for medical devices [1]. Collagen gels have also been used as matrices for drug delivery systems (DDSs) [2]. Collagen gels can be cross-linked by various methods to mechanically reinforce the gel, and also to control the rate of bioabsorption (degradation in vivo). However, the combination of mechanical strength and elasticity of the cross-linked collagen gels has so far not been adequate for applications in medical devices, especially in the clinical fields of orthopedics, cardiovascular surgery, and neurosurgery. It is clear that collagen gels of higher strength are needed [1,3–5]. Various methods for cross-linking, based on different principles, have been developed [1]. They can be broadly classified as i) natural cross-linking between hydroxyl lysine residues [1], ii) dehydrothermal cross-linking under vacuum at high temperatures, which forms intermolecular cross-links through condensation reactions, either esterification or amide formation [2,6], iii) treatment with glutaraldehyde (GTA) to effect crosslinking between two amino residues (\NH2) [1,2], iv) treatment with a carbodiimide (e.g., EDC) to crosslink between an amino residue (\NH2) and a carboxyl residue (\COOH) [1,7], v)
Abbreviations: EDC, 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide. ⁎ Corresponding author. Tel./fax: +81 72 254 9842. E-mail address: [email protected] (M. Hara). 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.03.047
other synthetic cross-linkers [1,2,8], vi) enzymatic reactions using transglutaminase [9], vii) other natural compounds as cross-linkers [10–12], viii) metal ions such as chromium(III), aluminum(III), zirconium(IV), iron(II) [13,14] and ix) radiation-based cross-linking through radical reactions (e.g., UV light, dye-mediated photooxidation, gamma rays, and electron-beam irradiation [1,5]). There is a continuing need for new cross-linking methods for collagen gels to reinforce the mechanical properties of the gel and avoid any toxicity due to residual reagents. From the various methods listed above, chemical cross-linking with GTA and with EDC has probably been the most widely used in tissue engineering and DDSs, although other cross-linkers of lower toxicity have emerged recently [1,2]. We therefore chose EDC as the cross-linker for the present study. We measured the dynamic viscoelasticities of collagen gels prepared and modified by four different methods: i) collagen gels crosslinked by EDC after preparation, ii) collagen gels cross-linked simultaneously with their preparation, iii) collagen gels irradiated with gamma rays after preparation, and iv) collagen gels formed directly from an acidic collagen solution by gamma cross-linking, as described previously [15]. We measured their mechanical properties using tensile testing and also the viscoelastic properties. From our results, we conclude that collagen gels with highly cross-linked thick collagen fibrils after heating showed a higher value of storage modulus (G′) and loss modulus (G″), compared with the previous report [15]. Using the experience obtained from our previous work, we succeeded in preparing collagen gels with extremely high mechanical strength by repeating the cycle of gel formation (fibril formation), cross-linking with EDC, and freeze-drying in which the heating was done only once after cross-linking in the first cycle. We also discuss the possible future uses of these materials.
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2. Materials and methods Materials and methods, including reagents, SEM observations, measurements of dynamic viscoelasticity, and tensile tests, were essentially as described in our previous paper [13]. We briefly describe the materials and methods. Porcine type I collagen solution (0.6% (w/v), pH 3.0, collagen BM, Nitta Gelatin, Osaka, Japan), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) (water-soluble carbodiimide hydrochloride 1030, Wako, Osaka, Japan), and other reagents were of a specific grade or analytical grade, as noted. A rheometer with an oscillatory rotating disk (Physica MCR301, Anton Paar GmbH, Graz, Austria) was used for dynamic rheological measurements. A tensile tester (Autograph AGS-50ND, Shimadzu, Kyoto, Japan) was used for the tensile tests. A freeze-dryer (FD-1, Eyela, Tokyo, Japan) and a deep freezer (Ultra Low, Sanyo, Tokyo, Japan) were used to lyophilize the gel. An electronic digital external micrometer (MCD130-25, Niigata Seiki, Sanjo, Japan) was used to measure the thickness of the collagen gel disks. An ion coater (IB-3, Eiko Engineering, Hitachinaka, Japan) and a scanning electron microscope (SEM; SM-100, Topcon, Tokyo, Japan) were used to characterize the gels. The preparation of the gels using cycled gel-formation is described below. Collagen gels were prepared by pH neutralization. Collagen (4 mL, 0.6% w/v), 0.8 mL of 10-fold concentrated phosphate-buffered saline (PBS), and 3.2 mL distilled water were mixed in a 6-well plate and incubated overnight at 37 °C to form the collagen gels. These were then incubated overnight in 10 mL PBS containing 50–125 mM EDC. The gels were washed well with distilled water to remove the salts. Each sample was placed on a plastic tray (35 mm in diameter, 20 mm in depth) and then wrapped with a thin plastic film, placed in a Styrofoam box (10 × 15 × 20 cm in length × width × depth) and frozen in a deep freezer at −85 °C overnight. The frozen sample was then lyophilized in a freeze-dryer to form the chemically cross-linked collagen sponge (microporous material). The sponge was immersed in the ice-cold collagen solution just after neutralization of the solution by addition of PBS. The solution promptly permeated into
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the sponge at 4 °C. The sample was then warmed and incubated at 37 °C to form the collagen gels inside the micropores in the sponge. The sample was incubated overnight in PBS containing 50 mM EDC for cross-linking. By repeating the cycle of collagen gel-formation followed by EDC cross-linking and lyophilization three times, we prepared what we refer to as the “3-cycle collagen gel 1” sample. A second mixture, consisting of 4 mL 0.6% collagen, 3.2 mL of distilled water, 0.8 mL of × 10 PBS(−), was prepared in a 6-well plate. After neutralization, the collagen gel was cross-linked with EDC and then heated for 30 min at 80 °C. The volume of the sample decreased significantly during the heating process. This gel is referred to as “1-cycle collagen gel”. The sample was lyophilized to form a sponge and was used to prepare a collagen gel followed by cross-linking with EDC twice. This gel is referred to as the “3-cycle collagen gel 2” in this paper. If the EDC-treatment was conducted only once, the gel was termed “2-cycle collagen gel 2”. Conditions for preparation of the various gels are shown with abbreviations in Tables 1 and 2 and in our previous paper [15]. All samples were prepared as described above and used for the rheological measurements in a wet state without additional lyophilization, and therefore are described as “gels”, not as “sponges”, in this paper and our previous paper [15]. Collagen gel disks (diameter, φ = 8–10 mm, thickness = 1–2 mm) were prepared and used for dynamic viscoelasticity measurements using a rheometer with an oscillatory rotating disk to obtain values for the storage modulus (G′) and the loss modulus (G″). Measurements were carried out with a deflection angle of 0.167–0.501 m rad and a range of angular frequency ω from 0.1 to 100 rad/s at ambient temperature (approximately 25 °C). Four samples (n = 4) were measured for each set of conditions, and the mean and standard deviation were calculated. The value of tangent δ (tan δ) was calculated as the ratio G ″/G′. This parameter reflects the thermal energy loss. The G′ value at 1 rad/s was assumed to be the G value. The number of network points per cubic meter (ν) was calculated using the formula: G ¼ νkT
Table 1 Summary of G′ and G″ values from the measurements of dynamic viscoelasticity. Abbreviations are Gel = gel-formation by neutralization of pH, H = heat-treatment for 30 min at 80 °C, 50E = the gel was treated with 50 mM EDC, and 125E = the gel was treated with 125 mM EDC. *Data recorded using conditions from Table 1 of the previous paper [13]. Curves, with standard deviations, are shown in Supplementary data 1–3. These values were slightly dependent on the range of rad/s (0.1–10) and frequency. Thus, values at the frequency of 1 rad/s are shown for comparison purposes in this table. Values of G′ and G″ are shown as means ± standard deviation (n = 4). If we compare pairs of unheated and heated gels, as analyzed by Student's t-test, no significant difference (p b 0.05) was found between the heated and unheated samples of any gel except between Gel/125E and Gel/125E + H (p b 0.01, as shown by #a). If we compare the heated sequential gel (Gel/125E + H) with both 2-cycled gel 2 and 3-cycled gel 2, values of viscoelastic parameters (G′ and G″) increased significantly with each cycle of gel formation. The superscript symbols with letters (#; p b 0.01, +; 0.01 ≤ p b 0.05) indicate significant differences between pairs of unheated and heated gels. Type of gel Abbreviation i) Uncross-linked collagen gel Gel*
G′ (Pa)
tan δ
G″ (Pa)
44.1 ± 31.3
ν (×1023, number/m3)
11.7 ± 8.8
0.265
0.107
ii) Sequentially EDC-cross-linked collagen gel (sequential gel) Gel/125E* 226 ± 70#a Gel/125E + H* 7010 ± 830#a,+c,+e,+g,#i
21 ± 4.4#b 288 ± 35#b,+d,#f,#h,#j
0.091 0.041
0.549 17.0
iii) 3-cycled collagen gel 1 (Gel/125E) × 3 times (Gel/125E) × 3 times + H
288 ± 35 168 ± 27
0.054 0.066
9.01 6.16
3730 ± 2060 2550 ± 517
iv) 2-cycled collagen gel 2 (Gel/50E) + H + (Gel/50E) (Gel/50E) + H + (Gel/50E) + H (Gel/125E) + H + (Gel/125E) (Gel/125E) + H + (Gel/125E) + H
25,400 16,200 20,400 26,100
± ± ± ±
13,700 2680 9780+c 11,700+e
842 657 944 1170
± ± ± ±
497 199 506+d 399#f
0.033 0.041 0.046 0.045
61.4 39.1 49.5 63.5
v) 3-cycled collagen gel 2 (Gel/50E) + H + (Gel/50E) × 2 times (Gel/50E) + H + (Gel/50E) × 2 times + H (Gel/125E) + H + (Gel/125E) × 2 times (Gel/125E) + H + (Gel/125E) × 2 times + H
18,000 16,800 40,200 38,500
± ± ± ±
6070 2800 18,000+g 6740#i
1300 765 3090 1920
± ± ± ±
416 94 1400#h 329#j
0.072 0.046 0.077 0.050
43.8 40.7 97.6 93.6
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Table 2 Summary of result from the tensile tests. Maximal values of the tensile stress (σ), tensile strain (ε) and corresponding Young's modulus (E) are shown as mean ± standard deviation (n = 7). *Data recorded using parameters from Table 1 of the previous paper [13]. Abbreviations are the same as in Table 1. There was no significant difference (p b 0.05) between the results from the sample with additional heating at the end (+H) and the corresponding unheated sample for each type of 3-cycled gel, as analyzed by Student's t-test. Comparing sequential gels, results were significantly different before and after the final heating step (+H) except for the values of ε for Gel/50E and Gel/50E + H, as described in Table 3 of the previous paper [13]. If we compare the results from the 3-cycled collagen gel 2 with those of sequential gels, the difference was significant. The superscript symbols with letters (#; p b 0.01, +; 0.01 ≤ p b 0.05) show that there were significant differences between the pair of unheated and heated gels. Type of gel Abbreviation
Tensile stress σ (MPa)
Tensile strain ε
Young's modulus E (MPa)
ii) Sequentially EDC-cross-linked gel (sequential gel) Gel/50E* Gel/50E + H*
0.010 ± 0.008#n 0.066 ± 0.019+a,#d,#n
4.37 ± 0.91 4.19 ± 1.28#b,#e
Gel/125E* Gel/125E + H*
0.004 ± 0.001#q 0.053 ± 0.011#h,#k,#q
3.34 ± 1.33 5.29 ± 1.50#i,#l
0.0047 0.013 0.097 0.0034 0.012 0.095
± ± ± ± ± ±
0.004#o,#p 0.0044#f,#o 0.035#c,#g,#p 0.0014#r,#s 0.005#j,#m,#r 0.034#s
v) 3-cycled collagen gel 2 (Gel/50E) + H + (Gel/50E) × 2 times (Gel/50E) + H + (Gel/50E) × 2 times + H (Gel/125E) + H + (Gel/125E) × 2 times (Gel/125E) + H + (Gel/125E) × 2 times + H
0.112 0.137 0.129 0.164
0.135 0.179 0.197 0.219
± ± ± ±
0.068#c 0.031#f,#g 0.069#j 0.051#m
± ± ± ±
0.050+a 0.035#d 0.040#h 0.063#k
where T is the absolute temperature (298 K) and k is the Boltzmann constant (1.38 × 10−23 J/K). Collagen gel disks, prepared as described, were pressed between two plastic sheets to a final size of φ = 8–10 mm, thickness = 1– 2 mm, and then subjected to tensile tests at ambient temperature (approximately 25 °C) using a tensile tester with an extension speed of 1 mm/min. Two pieces of sandpaper were placed on the surfaces of the sample folder to prevent slippage of the gels. Seven (n = 7) samples in each condition were prepared for the tensile test. The section area (A = thickness × diameter (mm 2)) of each sample was calculated from values measured using a digital micrometer. We did the tensile tests to obtain a stress (N)–strain (mm) curve. The original length of the sample in the tensile test (representing the distance between the upper and lower holders) was fixed at 1 mm. For a change in l of Δl, the tensile strain (ε) = Δl/l. The value of ε is thus the ratio between the original length and the change in length, and is therefore dimensionless. The tensile stress (σ) (MPa) = stress (N)/A (mm 2). Young's modulus (E) values were calculated from the maximal values of these parameters and the cross-sectional area of the sample gel, to give E = σ/ε. Various types of collagen gel, prepared in a 24-well plate, were washed with PBS and fixed for 3 h in PBS containing 2.5% (w/v) glutaraldehyde at 4 °C. The fixed samples were washed for 1 h in PBS, and then incubated successively for 20 min each in 30%, 50%, 70%, 80%, and 90% (v/v) ethanol at 4 °C. The samples were incubated for 20 min twice at ambient temperature to dehydrate them. The samples were then incubated for 20 min twice in a mixture of t-butanol and ethanol (1:1) to exchange the solvent, frozen in a freezer, and then dried under vacuum in a freeze-dryer. The dried samples were coated with platinum using an ion-coater and then observed using a scanning electron microscope (SEM). 3. Results We consider that both the thickness of the collagen fibrils and their 3D-density are important parameters for defining the viscoelastic properties of the collagen gels. Usually, these two parameters depend on the concentration of collagen used for the preparation of the gels. However, the acidic solution of 0.6% (w/v) collagen is already very viscous. It seems difficult to increase the concentration of the solution. We therefore invented a new method using repeated gel-formation inside the micropores of a collagen sponge. The collagen gels were prepared, cross-linked with EDC, washed and lyophilized, and then immersed in the collagen solution just before gel
1.26 1.14 1.06 1.07
± ± ± ±
0.32#b 0.28#e 0.21#i 0.23#l
formation to increase the 3D-density of the collagen fibrils in the gel. The number of cycles used is noted in Table 1, and described in the Materials and methods section, above. The shrinkage of the gels was quite distinct in the first heating cycle but less so in the second and third heating cycles. The three samples were used for SEM observations and rheological measurements of the dynamic viscoelasticity. The 2-cycled gel was white and non-transparent (data not shown). Fig. 1 shows SEM images of the 2-cycled gel prepared under various conditions. Collagen fibrils are observed in the gel treated with 50 mM EDC (A). These were twisted and fused to each other, forming thick bent fibrils with the inter-fibril space, reduced by additional heating (B). Thicker fused collagen fibrils were observed in the 2-cycled gels treated with 125 mM EDC before (C) and after the additional heating (D). The 3-cycled gel was also white and non-transparent, like the 2-cycled gel (data not shown). No significant change in size or appearance was observed after additional heating. Fig. 2 shows SEM images of the 3-cycled gel. The sample treated with 50 mM EDC showed a rough surface with fuzzy fibrils on the surface, as shown before (A) and after (B) heating. For the sample treated with 125 mM EDC, the fibrils were more densely packed and individual fibrils were difficult to observe, as in (C). The sample after heating showed a relatively smooth surface, as in (D). The interfibril spaces in (C) and (D) are believed to be small because of the densely packed collagen fibrils. The G′ and G″ values obtained from the measurements of dynamic viscoelasticity are summarized in Table 1. As mentioned previously, the gel with the highest values of G′ and G″ was the 3-cycled gel ((Gel/125E) + H + (Gel/125E) × 2 times). The second highest was that with additional heating ((Gel/125E) + H + (Gel/125E) × 2 times + H). The third highest was the 2-cycled gel ((Gel/125E) + H + (Gel/125E) + H). All of the 3-cycled gel 2 samples had higher values of G′ and G″ (G′ > 16,000 Pa, G″ > 700 Pa) than those of the 3-cycled gel 1 (G′ b 4000 Pa, G″ b 300 Pa) without heating after their first cross-linking ((Gel/125E) × 3 times, and (Gel/125E) × 3 times + H). The gel cross-linked once with EDC without heating (Gel/125E) showed higher values (G′; 226 Pa, G″; 21 Pa) than the uncrosslinked gel (Gel) (G′; 44.1 Pa, G″; 11.7 Pa). However those values were far lower than those (G′; 7010 Pa, G″; 288 Pa) of the sample with heating (Gel/125E + H). As described in our previous paper, the first heating made the gel shrink, causing the collagen fibrils to bend, twist, and fuse to each
H. Mori et al. / Materials Science and Engineering C 33 (2013) 3230–3236
A
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B
2.0 µm
C
2.0 µm
D
2.0 µm
2.0 µm
Fig. 1. SEM images of the 2-cycled gels, (Gel/50E) + H + (Gel/50E) (A), (Gel/50E) + H + (Gel/50E) + H (B), (Gel/125E) + H + (Gel/125E) (C), and (Gel/125E) + H + (Gel/ 125E) + H (D). The images show the surface of the gels.
other [15]. We conclude that these changes brought about by heating contributed to the increase of G′ and G″ values, in addition to the cross-linking. We also need to understand the effect of the final heating on the cycled and non-cycled gels. The gel, cross-linked once with 125 mM EDC, showed a distinct increase in G′ value, from 226 to 7010, by
heating. However, a 2-cycled gel cross-linked with 125 mM EDC showed only a slight increase in the G′ value, from 20,400 to 26,100, after the final heating while the 3-cycled gel cross-linked with 125 mM EDC showed a decrease in G′ value, from 40,200 to 38,500. Thus, it appears that the strengthening effect brought about by heating after cross-linking is a distinct process after the first
Fig. 2. SEM images of the 3-cycled gels, (Gel/50E) + H + (Gel/50E) × 2 times (A), (Gel/50E) + H + (Gel/50E) × 2 times + H (B), (Gel/125E) + H + (Gel/125E) × 2 times (C), and (Gel/125E) + H + (Gel/125E) × 2 times + H (D). The images show the surface of the gels.
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in the same direction, parallel to the extension, and therefore show low tensile stress. In the moderately increasing phase, bundles of the thick collagen fibrils are mechanically strained, showing higher tensile stress. We expect that the shapes of the curves for the 3-cycled gel 2 shown in Fig. 3 of this paper can be described similarly. A network of thick twisted collagen fibrils (diagram B in Fig. 7 of Ref. [15]) and that of fibrils of intermediate thickness (diagram C in Fig. 7 of Ref. [15]) become interpenetrated and entangled with each other. Both types of fibrils become highly crosslinked with EDC. Thus, bundles of thick collagen fibrils are mechanically strained, showing higher tensile stress than the moderately-increasing phase. The rapidly increasing phase was diminished because almost all the heat-denatured collagen (gelatin) was already crosslinked with EDC. Data obtained in the tensile test are summarized in Table 2. The values of Young's modulus were somewhat variable, falling within the range of the 1.3-fold change in the sample treated with both 50 mM and 125 mM EDC with and without heating. There was no significant difference (p b 0.05) between the heated sample and unheated one for each set of conditions, as analyzed by Student's t-test. From the data in Table 2, we conclude that all four types of 3-cycled gel 2 were mechanically strong. The tensile strain was in the range from 1.26 to 1.06, so we also conclude that the 3-cycled gel 2 samples have low tensile strain.
cross-linking, but becomes less apparent after further cross-linking cycles. With 50 mM EDC, even the 2-cycled gel showed a decrease in the G′ value with heating, from 25,400 to 16,200, suggesting the possibility that uncross-linked collagen had been partially dissolved. There was no significant difference (p b 0.05) between the heated sample and unheated one for each set of conditions in Table 1, as analyzed by Student's t-test, except the comparison between Gel/125E and Gel/125E + H. Thus, we conclude that the final heating process did not greatly alter the rheological properties. Examples of the tensile test data (tensile strain–tensile stress curve) for the 3-cycled gel 2 are shown in Fig. 3. The shapes of the curves for gels prepared using four different conditions (Fig. 3(A)–(D)) were similar to each other. The initial increase in the tensile stress was almost linear and then the stress gradually reached a plateau. The maximal tensile stress at a final breaking point was variable, and in the range 0.05–0.25. The strain at the maximal tensile stress was typically small (ε b 1.5) in the sample treated with 125 mM EDC, suggesting significant stiffness of the sample. We measured the tensile test of the sequentially EDC-cross-linked collagen gel, as shown in our previous paper [15]. The shapes of the curves for the unheated sequentially EDC cross-linked collagen gel consist of three phases – a lag phase, a moderately increasing phase, and a plateau – while another rapidly-increasing phase was present for the heated gel, as shown in Fig. 3 of our previous paper [15]. By comparison with these data, the curves for the present 3-cycled gels clearly have neither a lag phase nor a rapidly-increasing phase. Why the shapes of the curves are so different between the cycled gel 2 in this paper and the sequential gel in the previous paper [15] is not easy to answer but seems to relate to the mechanism of deformation of the gel in the tensile test. We believe that the rapidly-increasing phase in the heated sample of sequential gels is derived from the heat-denatured collagen (gelatin), as described in the previous paper [15]. We now consider the shapes of the curves in the sequential gel using the previous arguments [15]. The initial lag phase in the unheated sequential gels corresponds to the process in which the cross-linked collagen fibrils are mechanically forced to be oriented
4. Discussion We succeeded in preparing a mechanically durable collagen gel using successive cycles of gel formation by neutralization, cross-linking with EDC, lyophilization, and heating. We determined that the 3-cycled collagen gel 2, processed with 125 mM EDC but without heating, was mechanically very strong (storage modulus G′; 40,200 Pa, loss modulus G″; 3090 Pa, Young's modulus E; 0.197 MPa, tensile stress σ; 0.129 MPa, tensile strain ε; 1.06) and shows promise as a scaffold for tissue engineering research. This 3-cycled gel 2 was mechanically the strongest collagen gel we have prepared. However,
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Fig. 3. Examples of the tensile stress–tensile strain curves obtained in the tensile test of (Gel/50E) + H + (Gel/50E) × 2 times (A), (Gel/50E) + H + (Gel/50E) × 2 times + H (B), (Gel/125E) + H + (Gel/125E) × 2 times (C), and (Gel/125E) + H + (Gel/125E) × 2 times + H (D). Different line types show different examples of results using the same measurement conditions.
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the 3-cycled gel 2 was stiff and not tensible because the tensile strain was only 1.06. We considered why the cycled gels were mechanically much stronger than conventional collagen gels crosslinked once with EDC but show only moderate increases in the values of their Young's modulus. However, we currently have no clear answer to this question. It is possible that a tiny crack formed in the sample, causing it to tear during the tensile test, and so reduce the maximal tensile stress, finally resulting in a moderate value of Young's modulus. As also discussed in our earlier paper [15], a mechanically strong scaffold is sometimes necessary, in the tissue engineering of bone, cartilage, tendon, ligament, muscles, and other tissues. In particular, if we culture chondrocytes to repair injured cartilage, mechanically strong collagen gels or collagen sponges are definitely necessary [16]. The mechanical properties of tissues are determined by the density and orientation of the collagen fibrils, the extent of covalent cross-linking between the collagen molecules, and the contribution of other extracellular matrix (ECM) components, such as elastin, glycosaminoglycans, and proteoglycans. Collagen fibrils are in relatively random orientations in the dermis and cartilage, but are highly oriented, unidirectionally, in tendons, ligaments, and trabecular bone. Cross-linked networks of collagen and elastin are also contained in the laminated layers of the walls of blood vessels, providing their tensibility and elasticity. We prepared the 3-cycled collagen gel 2 with good mechanical strength, although the orientation of the collagen fibrils was not controlled in this research. We consider that a scaffold for cartilage is a candidate use of this gel, in tissue engineering [16]. However, the pore size appears to be small and will not permit cells to proliferate into the scaffold material. This can be advantageous or disadvantageous, depending on the application. This problem should be addressed by additional steps of lyophilization or some other method in the future. We used 0.6% (w/v) porcine atelocollagen that is commercially available. We did not try to concentrate the collagen solution further in this study. If the viscosity of the solution is too high, it is difficult to mix the solution with PBS for neutralization. It has been reported that inter-molecular interactions are present in very concentrated solutions of collagens [17]; liquid crystal-like nematic phases and precholesteric phases appear in 2% (w/v) and 4% (w/v) solutions, respectively. These intermolecular interactions may contribute to the mechanical properties of the collagen gel by affecting the thickness and orientation of the collagen fibrils. These issues remain to be further characterized. In the preparation of the gel with the highest G′ value ((Gel/ 125E) + H + (Gel/125E) × 2 times), the collagen gel was made into a porous collagen sponge by the first freeze-drying, with crystals of ice acting as porogens. Then, the second fibril-formation occurred in the pores (inner spaces) of the sponge thus forming the hierarchical structure of the collagen fibrils. The first fibrils to be formed were partially denatured into thick twisted fibrils or were fused together by the heating, while the second and third fibrils were not. The denatured thick collagen fibrils and undenatured fibrils show different mechanical properties, as described in our previous paper [15]. We assume that we made two collagen gels, with differing mechanical properties, even though we used the same raw material, collagen. It has also been reported that the pore size of the collagen sponge is an important factor governing cell migration into the sponge. If the pore size is too large, the surface area is small. If the pore size is too small, the cells find it difficult to migrate into the sponge. O'Brien and colleagues tested in detail the effects of various conditions (e.g., collagen concentration, freezing rate, temperature) on the preparation of the sponge by freeze-drying [18–20] and concluded that a mean pore size of 120 μm was the most suitable for the migration of cells into a collagen–glycosaminoglycan sponge [21]. We have
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not yet measured the pore size or tested the migration of cells into the cycled gels prepared in this study but this will be a subject for future investigations. The mechanical properties – especially the tensile strength – of the collagen gel fabricated also depend on the orientation of the collagen fibrils. It has also been reported that a collagen gel block, once prepared, can be compressed mechanically to a thin film, and then sequentially cross-linked with EDC to form an EDC-crosslinked collagen gel sheet [22–24]. The orientation of the collagen fibrils along the z-axis was not controlled, but neither was the orientation along the x- or y-axis. This method can effectively increase the three-dimensional density of the collagen fibrils. Sheets or films consisting of collagen gel are useful as structural biomaterials. They can be cut with scissors, laminated to form multi-layers [25], or rolled up to form concentric lamellae, like construction paper [22–24,26]. Zeugolis et al. prepared extruded collagen fibers in which collagen fibrils were oriented [27,28], and then treated them by various crosslinking methods, as listed in Table 2 of Ref. [28]. Tanaka et al. reported the preparation of collagen gel sheets in which the collagen fibrils were highly oriented, using an oriented flow casting method [26]. We have not yet tried to control the orientation of collagen fibrils in the present work but this will also be an interesting theme for future investigations. 5. Conclusions We have described the preparation of mechanically durable collagen gels (G′; 40,200 ± 18,000 Pa, G″; 3090 ± 14,000 Pa) by repeating the gel-formation/EDC cross-linking/lyophilization cycle. The gels prepared by various methods including cross-linking and heating, need to be further characterized in terms of the technical issues of cell adhesion properties, biocompatibility, and bioabsorptivity (non-toxic biodegradability in vivo) before their possible use in medical devices. Acknowledgments We thank Dr. Shiho Suzuki, and Shinichi Kitamura in Osaka Prefecture University for their support in the measurements of dynamic viscoelasticity, and to Ms. Hiroko Machida of AIST-Kansai for SEM observations. Appendix A. Supplementary data Supplementary data to this article can be found online at http:// dx.doi.org/10.1016/j.msec.2013.03.047. References [1] S.F. Badylak, in: A. Atala, R.P. Lanza (Eds.), Methods of Tissue Engineering, Academic Press, London, 2002, pp. 503–514. [2] W. Friess, Eur. J. Pharm. Biopharm. 45 (1998) 113. [3] I.V. Yannas, J. Macromol. Sci., Rev. Macromol. Chem. C7 (1972) 49B. [4] I.V. Yannas, Clin. Mater. 9 (1992) 179. [5] M. Hara, J. Oral Tissue Eng. 3 (2006) 118. [6] M.G. Haugh, M.J. Jaasma, F.J. O'Brien, J. Biomed. Mater. Res. 89A (2009) 363. [7] S. Yunoki, K. Mori, T. Suzuki, N. Nagai, M. Munekata, J. Mater. Sci. Mater. Med. 18 (2007) 1369. [8] E.C. Collin, S. Grad, D.I. Zeugolis, C.S. Vinatier, J.R. Clouet, J.J. Guicheux, P. Weiss, M. Alini, A.S. Pandit, Biomaterials 32 (2011) 2862. [9] J.M. Orban, L.B. Wilson, J.A. Kofroth, M.S. El-Kurdi, T.M. Maul, D.A. Vorp, J. Biomed. Mater. Res. 68A (2004) 756. [10] D.I. Zeugolis, R.G. Paul, G. Attenburrow, Mater. Sci. Eng. C30 (2010) 190. [11] F.-L. Mi, C.-T. Huang, Y.-L. Chiu, M.-C. Chen, H.-F. Liang, H.-W. Sung, J. Biomed. Mater. Res. 83A (2007) 667. [12] H.-W. Sung, W.-H. Chang, C.-Y. Ma, M.-H. Lee, J. Biomed. Mater. Res. 64A (2003) 427. [13] A.D. Covington, Chem. Soc. Rev. (1997) 111. [14] B. Wu, C. Mu, G. Zhang, W. Lin, Langmuir 25 (2009) 11905. [15] H. Mori, K. Shimizu, M. Hara, Mater. Sci. Eng. C32 (2012) 2007.
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[16] J.M. MacPherson, R. Tubo, in: R.P. Lanza, R. Langer, J. Vacanti (Eds.), Principles of Tissue Engineering, 2nd edition, Academic Press, London, 2000, pp. 697–709. [17] L. Besseau, B. Coulomb, C. Lebreton-Decoster, M.-M. Giraud-Guille, Biomaterials 23 (2002) 27. [18] F.J. O'Brien, B.A. Harley, I.V. Yannas, L.J. Gibson, Biomaterials 26 (2005) 433. [19] C.M. Tierney, M.G. Haugh, J. Liedl, F. Mulcahyy, B. Hayes, F.J. O'Brien, J. Mech. Behav. Biomed. Mater. 2 (2009) 202. [20] F.J. O'Brien, B.A. Harleyb, I.V. Yannas, L. Gibson, Biomaterials 25 (2004) 1077. [21] C.M. Murphy, M.G. Haugh, F.J. O'Brien, Biomaterials 31 (2010) 461. [22] R.A. Brown, M. Wiseman, C.-B. Chuo, U. Cheema, S.N. Nazhat, Adv. Funct. Mater. 15 (2005) 1762.
[23] M. Bitar, V. Salih, R.A. Brown, S.N. Nazhat, J. Mater. Sci. Mater. Med. 18 (2007) 237. [24] E.A.A. Neel, U.C. Jonathan, C. Knowles, R.A. Brown, S.N. Nazhat, Soft Matter 2 (2006) 986. [25] M. Hara, A. Yamaki, J. Miyake, Mater. Sci. Eng. C17 (2001) 107. [26] Y. Tanaka, K. Baba, T.J. Duncan, A. Kubota, T. Asahi, A.J. Quantock, M. Yamato, T. Okano, K. Nishida, Biomaterials 32 (2011) 3358. [27] D.I. Zeugolis, R.G. Paul, G. Attenburrow, J. Biomed. Mater. Res. 86A (2008) 892. [28] D.I. Zeugolis, G.R. Paul, G. Attenburrow, J. Biomed. Mater. Res. 89A (2009) 895.