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Proteoglycans and Glycosaminoglycans Improve Toughness of Biocompatible Double Network Hydrogels Yu Zhao, Tasuku Nakajima, Jing Jing Yang, Takayuki Kurokawa, Jian Liu, Jishun Lu, Shuji Mizumoto, Kazuyuki Sugahara, Nobuto Kitamura, Kazunori Yasuda, A. U. D. Daniels, and Jian Ping Gong* Hydrogels made from biopolymers, such as hyaluronan and collagen, are widely used in biomedical applications as cell scaffolds and artificial tissues.[1–3] These biopolymer hydrogels, which are either chemically or physically cross-linked, are usually very soft with nominal tensile moduli ranging from 10 to 104 Pa.[4] Thus, like conventional hydrogels made from synthetic single polymers, these biopolymer hydrogels are much less stiff and strong than load-bearing biological tissues. For example, cartilage has a nominal modulus of 0.1–1.0 MPa, a compressive fracture stress of several megapascals and a tearing energy of more than 1000 J m–2,[5] while hyaluronan gels typically show a fracture stress of only 105 Pa.[6] Development of biopolymer-based hydrogels with mechanical performance comparable to that of load-bearing soft tissues would greatly broaden the application of these materials as substitutes for tissues such as cartilage. The double-network (DN) principle provides a general approach to developing hydrogels with cartilage-like robustness but also high water content (ca. 90 wt%).[7] One family of tough DN hydrogels consists of poly(2-acrylamido-2-methylpropanesulfonic acid) (PAMPS) as the first network and poly(acrylamide) (PAAm) as the second network. They have tensile fracture stress and tearing fracture energy as high as 2 MPa and 4000 J m–2, respectively, comparable to or even superior to cartilage.[8–10] Many studies have shown that this DN concept is general and does not depend on specific chemical interactions between the two networks.[11,12] Thus, the DN concept should be applicable to developing tough hydrogels with biochemical functionality using combinations of various biopolymers and biocompatible polymers.
The challenge is how to synthesize the specific structure required by the DN hydrogels using biopolymers and biocompatible polymers. Synthetic polymer DN gels achieve toughness by having a topological structure very different from a conventional interpenetrating network (IPN). It consists of two interpenetrating networks that have contrasting physical structures and properties: a densely cross-linked and rigid (and thus relatively brittle) network (skeleton) of low concentration and a sparsely cross-linked soft and ductile network of high concentration.[7,8] Specifically, the molar mass of the second network is optimally 7–20 times that of the first network. For synthetic polymers, this contrasting structure is formed by a two-step sequential polymerization process: first one synthesizes a polyelectrolyte network (e.g., PAMPS) as the rigid skeleton, and then one synthesizes within this structure a neutral network (e.g., PAAm) as the soft and ductile network. Following this approach, one strategy to synthesize biopolymer-based DN gel would be to use a biopolymer as the first network and a neutral biocompatible polymer as the second network. This is because biopolymers, such as proteoglycans or glycosaminoglycans, are highly charged polyelectrolytes. Following this approach, tough DN hydrogels using chemically modified methacrylate chondroitin sulfate as the first network, and biocompatible poly(N,N-dimethyl acrylamide) (PDMAAm) as the second network have been developed recently.[13] However, this approach requires chemical cross-linking of the biopolymer, which is a complicated process. Another disadvantage is that chemical cross-linking might cause unfavorable changes in the biochemical properties of the biopolymers.
Prof. J. P. Gong Laboratory of Soft & Wet Matter Faculty of Advanced Life Science Hokkaido University Kita-10-Nishi-8, Kita-ku, Sapporo, 060-0810, Japan E-mail: [email protected] Y. Zhao Laboratory of Soft & Wet Matter Graduate School of Life Science Hokkaido University Kita-10-Nishi-8, Kita-ku, Sapporo, 060-0810, Japan Dr. T. Nakajima, Dr. J. J. Yang, Dr. T. Kurokawa Laboratory of Soft & Wet Matter Faculty of Advanced Life Science Hokkaido University Kita-10-Nishi-8, Kita-ku, Sapporo, 060-0810, Japan
J. Liu Laboratory of Soft & Wet Matter Graduate School of Science Hokkaido University Kita-10-Nishi-8, Kita-ku, Sapporo, 060-0810, Japan J. Lu, S. Mizumoto, Prof. K. Sugahara Laboratory of Proteoglycan Signaling and Therapeutics Faculty of Advanced Life Science Hokkaido University Kita-21-Nishi-11, Kita-ku, Sapporo, 001-0021, Japan Dr. N. Kitamura, Prof. K. Yasuda Department of Sports Medicine and Joint Surgery Graduate School of Medicine Hokkaido University Kita-15-Nishi-7, Kita-ku, Sapporo, 060-0821, Japan Prof. A. U. D. Daniels Biomechanics & Calorimetry Center Basel Faculty of Medicine University of Basel c/o Biozentrum, Klingelbergstrasse 50-70, 4056 Basel, Switzerland
DOI: 10.1002/adma.201303387
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Here, we present a more general approach 2nd network twork to synthesizing biopolymer-containing tough 2nd monomer onom st ne 1 network DN hydrogels. It is based on the so-called Molecular Swelling in 2nd molecular stent method that we developed stent nt 2nd solution gelation recently.[14] It can be used to combine a chemically inert, neutral, biocompatible polymer network with any polyelectrolyte biopolymer to form a tough DN gel without any chemical St gel modification of the biopolymer. A hydrogel (as-prepared, formed from a neutral synthetic polymer St gel St-DN gel coil state) swells only modestly since the polymer seg(Swollen state) ments take on a coiled conformation in water. To use a neutral polymer as the first network Molecular stent 1st network , 2nd network of a DN gel, one needs to make the network Chondroitin sulfate proteoglycans (PG), rigid, as is the case with a polyelectrolyte. To Sodium hyaluronate (HA), Monomer do this, our strategy is to use a biopolymer Chondroitin sulfate (CS) N,N-Dimethylacrylamide (DMAAm) as a molecular stent, taking advantage of its Proteoglycan polyelectrolyte nature. As shown in Figure 1, DMAAm we add the biopolymer to the precursor monChondroitin sulfate omer solution of the first neutral network. Hyaluronan After the gel formation by polymerization, Crosslinker N, N’-methylenebis(acrylamide) (MBAA) the biopolymer trapped in the neutral network exerts extra ionic osmotic pressure on the network. As a result, the neutral gel MBAA swells far more substantially, and its polymer Proteoglycan aggregate strands become extended and rigid and consequently relatively brittle, as happens in Figure 1. Schematic illustration of how polyelectrolyte biopolymers facilitate formation of the polyelectrolyte gels. After that, we can intro- contrasting DN structure from a neutral biocompatible polymer. A hydrophilic hydrogel with duce a large amount of a second neutral a charge-neutral chain exhibits only modest swelling, with the polymer strands remaining in network to form the contrasting topological coiled conformation in water. When a polyelectrolyte biopolymer is entrapped in the structure it structure required for a tough (non-brittle) exerts an extra ionic osmotic pressure on the neutral network. As a result, the neutral network DN gel. The advantages of this method are swells, and its polymer strands (pink) are fully stretched, taking on an extended conformation that 1) we can introduce the biopolymers as is the case in polyelectrolyte gels. Thus the polyelectrolyte biopolymer acts as a molecular stent. Owing to the substantial swelling of the first network, an excess amount of the second without any chemical modification; 2) difmonomer can be incorporated afterward, and the second neutral network with high concentraferent biopolymers can be introduced into tion can be synthesized. As a result, the contrasting topological structure of a tough DN gel is the DN gel simultaneously with any desired formed, and simultaneously a biopolymer has been loaded into the tough DN gels. composition, mimicking the structure of soft tissues; 3) any neutral biocompatible polymer can be used as We turn now to the stiffness of St-SN gels. The high molethe first network of a biopolymer-stented DN gel. cular weights of biopolymers make their solutions extremely Using the above-mentioned approach, we have successfully viscous. As a result, the maximum biopolymer concentration developed a series of tough DN hydrogels from biopolymers of Cstent in the precursor solution of the hydrogels was limited chondroitin sulfate proteoglycans (PGs), one of the major comto 5 wt%. Above this concentration, it was difficult to prepare ponents of cartilage, and their glycosaminoglycan components, homogeneous solutions. Figure 2b shows the volume change chondroitin sulfate (CS), and sodium hyaluronate (HA), in qv (=Vswell/Vprep) of the PDMAAm St-SN gels relative to the combination with a neutral biocompatible polymer, poly(N,Nas-prepared state for samples containing the three biopolydimethylacrylamide) (PDMAAm).[15–18] These hydrogels, which mers. A substantial increase in qv is observed with Cstent. The we call St-DN gels, exhibit mechanical properties comparable to nominal tensile modulus E of the gels also increases with the standard all-synthetic PAMPS/PAAm DN hydrogels[9,10] but are incorporation of biopolymers (Figure 2c). The tensile stiffness in addition potentially bioactive. The molecular weight and charge density of biopolymers used Table 1. Molecular weight, charge density, and origin of biopolymers in this work are shown in Table 1. The molecular stent effect of used in this work. biopolymers is demonstrated in Figure 2a. The photos show (a, left) a highly swollen PDMAAm single-network hydrogel conBiopolymers Molecular weight Charge density Sources taining a PG stent, which we call St-SN gel, and (a, right) the [Da] [mol g–1] less-swollen PDMAAm gel without the PG stent. The gel con2.98 × 10−3 salmon nasal cartilage PGs 2.71 × 106 taining PG as a molecular stent was selectively stained by Alcian 6 2.63 × 10−3 HA (1.8–2) × 10 Streptococcus equi subsp. blue, while the PDMAAm gel without PG could not be stained zooepidemicus by Alcian blue. These results demonstrate that entrapping of 5 −3 2.42 × 10 squid cartilage CS (0.7–3) × 10 PGs leads to enhanced swelling of the neutral hydrogel.
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Figure 2. Effects of biopolymers on swelling of the first network composed of the neutral polymer PDMAAm. a) The SN gel containing PGs (St(PGs)SN gel) (left) and without PGs (right) after dyeing with Alcian blue. The blue color confirms the presence of the PGs and the increased volume of the St(PGs)-SN gel demonstrates the molecular stent effect created. The swelling degree qv (b), nominal elastic modulus in tension E (c), and relative modulus En (d) of St-SN gels with various concentrations of biopolymers Cstent. qv = Vswell/Vprep, En = qvEswell/Eprep, where Vswell (Eswell) and Vprep (Eprep) are the volume (modulus) of the swollen gel and the as-prepared gel, respectively. Cstent is the feed-in value of the biopolymer in the precursor solution of the gel. Sample formulation: St-2–2.
(e.g., nominal elastic modulus E) of a bulk hydrogel is correlated to both the density (number per unit volume) of elastically effective polymer strands ve and the energy of a single strand echain, E = 3veechain.[19] The biopolymers were physically entrapped in the gel and thus they did not make a significant contribution to the stiffness, while the ve of the PDMAAm strands is inversely proportional to the swelling ratio qv. Thus, the increase of E means that echain increases with the swelling, surpassing the decreasing effect of ve. As echain is also related to the stiffness of the strand, the stiffness of the PDMAAm strand in the swollen St-SN gels relative to the as-prepared state can be estimated from the normalized nominal tensile modulus En of the St-SN gels from the following equation (in which E = Eswell), En =
Vswell Vprep
E swell E prep
= qv
E E prep
(1)
As shown in Figure 2d, En increases with the biopolymer concentration Cstent. These results demonstrate that the stiffness of each neutral strand increases with the swelling due to the ionically driven osmotic pressure πion of the biopolymers. Since πion = cionRT, where cion is the concentration of the dissociated counter ions of the biopolymers, we further investigated the molecular stent effect of biopolymers in terms of their charge concentration (Figure S2, Supporting Information). The charge 438
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concentrations of biopolymers in the precursor solutions was less than 0.07 M, about 1/10 that of the synthetic polymer stent used in our previous work.[14] Consistent with this, the relative volume change qv and modulus change En were also about 1/10 those obtained using synthetic molecular stents.[14] Among the three biopolymers, CS showed the weakest stent effect at the same polymer weight concentration Cstent (Figure 2) or charge concentration (Figure S2). This is because of a higher leakage rate of CS from the gel. Owing to its relatively low molecular weight, about 50% of CS had leaked from the gel after one day, at which time the specimens were measured (Figure S3, Supporting Information), so the actual concentration of CS was only 50% in the St(CS)-SN gels. Owing to their much higher molecular weights, negligible leakage was observed in the St(PGs)- and St(HA)-SN gels. The above results are in agreement with previous studies and confirm again that the swelling effect induced by the molecular stent is universal and independent of the chemical species of the polyelectrolyte used. We also analyzed and optimized the St-DN gel stress–strain behavior. The DN gels containing biopolymers (St-DN gels) were synthesized by immersing the St-SN gels in the precursor solution of the second network for one day and performing the radical polymerization again. The modest swelling of the St-SN gels due to the low charge concentration of the biopolymers is
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Table 2. Sample codes, compositions, and mechanical properties of St-DN gels containing biopolymers. Sample code
Formulation Cstent [wt%]
x1–y1/x2–y2
Mechanical propertiesa)
True composition a)
b)
Stent conc. [wt%]
Polymer conc. [wt%]
Water conc. [wt%]
[mol/mol]
E [MPa]
σf [MPa]
C2nd/C1sta)
εf [%]
Wf [J m–3]
St(0)-DN
0
1–2/4–0.02
0
8.8
91.2
2.99
0.07
0.19
125
0.15
St(PGs)-DN1
1
1–2/4–0.02
0.14
10.26
89.6
6.27
0.11
1.5
174
0.85
St(PGs)-DN2
3
1–2/4–0.02
0.33
10.98
86.69
8.46
0.13
0.91
195
0.87
St(PGs)-DN3
5
1–4/4–0.02
0.82
12.18
87
6.98
0.33
0.96
473
3.7
St(HA)-DN1
1
1–2/4–0.02
0.16
10.44
89.4
5.78
0.1
1.15
197
1.16
St(HA)-DN2
2
1–2/4–0.02
0.55
10.74
88.7
8.60
0.14
1.06
168
0.8
St(HA)-DN3
3
1–2/4–0.02
0.58
11.14
88.2
10.70
0.19
0.73
173
0.75
St(CS)-DN1
1
1–2/4–0.02
0.09
10.11
89.8
4.91
0.09
1.36
185
0.78
St(CS)-DN2
3
1–2/4–0.02
0.18
10.42
89.4
8.09
0.11
1.2
181
0.86
St(CS)-DN3
5
1–2/4–0.02
0.28
10.62
89.1
8.73
0.12
1.04
174
0.85
St(CS)-DN4
8
0.9–4/4–0.02
0.47
11.1
88.43
10.07
0.3
0.71
426
2.26
a)In
x1–y1/x2–y2,x and y denote the DMAAm monomer concentration [M] and the MBAA crosslinker density [mol%] with respect to the monomer, respectively. The subscripts1 and 2 denote the first and second networks, respectively. C2nd/C1st is the molar ratio of DMAAm units of the second network to that of the first network. E, σf, εf, and Wf are nominal tensile elastic modulus, fracture stress, fracture strain, and work of extension at fracture of St DN gels, respectively; b)True weight contents of biopolymers, taking into consideration the leakage over time of biopolymers.
unfavorable to inducing an excess amount of the second network relative to the first network, as required to form the contrasting structures that give DN gels their exceptional mechanical properties. In order to overcome this problem, the DMAAm monomer concentration of the second network was doubled (4 M) in comparison with the previous study.[14] The formulation, the true composition, and the mechanical properties of the St-DN gels are shown in Table 2. These samples contained about 86–91 wt% water, and the molar ratio of DMAAm of the second network to that of the first network, C2nd/C1st, showed an increasing trend with the concentration of molecular stent, reaching values of 3–11. The composition data shown in Table 2 indicate that the contrasting DN structure was indeed formed in the St-DN gels. Single-network St-SN gels, packed with linear biopolymers, were relatively low in stiffness and strength, as shown by their tensile stress–strain curves in Figure 3a. Also, stress increased monotonically with the strain, and then the gels fractured abruptly with little or no plastic deformation (brittle behavior). The fracture stress and fracture strain were less than 0.1 MPa and 60%, respectively. In contrast, the double-network St-DN gels showed quite different stress–strain behavior (Figure 3b). The fracture stress and fracture strain increased dramatically after formation of the second network. Furthermore, the fracture stress and fracture strain of St-DN gels were remarkably high compared to those made without biopolymers (Table 2). Figure 3c shows the images of a St(PGs)-DN gel withstanding 500% elongation. Also all St-DN gels showed yielding (plastic deformation) behavior before failure, a characteristic feature of tough DN gels, while no yielding was observed for specimens not containing biopolymers. Quantitatively, the work of extension at tensile fracture, a parameter that expresses one aspect of toughness of gels, increased dramatically in the case of St-DN gels (Table 2). Studies on the toughening mechanism of DN gels have shown that eventual yielding (plastic deformation)
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under tension is due to the internal fracture of the first network. Bond breakage in this network dissipates energy that would otherwise result in crack propagation, and thus imparts toughness to the DN gels.[8,11,20] Thus, the yielding behavior of the St-DN gels corresponds to the internal fracture of the PDMAAm first network and the toughening mechanism of the St-DN gels is the same as that of a conventional PAMPS/PAAm DN gel that consists of the polyelectrolyte PAMPS as the first network and the neutral polymer PAAm as the second network. It should be emphasized that the linear biopolymers, physically entrapped in the double network, only serve as molecular stents that facilitate formation of the rigid, extended structure of the first network. They do not transfer force between the two networks during deformation and thus play no role mechanically in the enhancement of the double network gels. Without the biopolymers, the contrasting topological structures could not be formed in neutral double networks, and in such a case a DN gel behaves like a conventional IPN gel, showing little or no yield under tension, and thus elastic deformation followed by brittle fracture. First-generation St-DN biopolymer gels have some limitations. Compared to standard DN gels, and St-DN gels made using a synthetic polymer as a molecular stent (St-DN synthetic gels), the fracture strain of the St-PDMAAm/PDMAAm DN gels made using biopolymer stents was 2–3 times smaller.[11,14] Furthermore, the stress–strain curves of St-DN gels showed a short flat region after yielding, but then an upward slope after that (Figure 3b). Similar behaviors were also observed in the StPAAm/PAAm DN gels, when PAAm was used as both the first and the second network with biopolymers as molecular stents (Figure S4, Supporting Information). We consider that this behavior stems from the different composition ratio C2nd/C1st of the second network to the first network in these cases. For a standard DN gel and St-DN synthetic gel, optimized C2nd/C1st was around 20. In the case of the St-DN biopolymer
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Figure 3. Effect of biopolymers on the mechanical behavior of SN gels and DN gels composed of the neutral polymer PDMAAm. a,b) Tensile stress–strain curves of single-network gels (St-SN gels) (a) and doublenetwork gels (St-DN gels) (b). c) Photographs demonstrating how a St-DN gel sustains a large strain of 500%. The sample codes of the St-DN gels used in (b) are (䊏) St(0)-DN, (ⵧ) St(PGs)-DN3, (䉭) St(HA)-DN3, (x) St(CS)-DN4. The sample code of the sample used in (c) is St(PGs)DN3. The compositions are shown in Table 1. The St-SN gels in (a) have first-network compositions corresponding to those of the St-DN gels in (b). Note that different scales are used in (a) and (b).
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gels, C2nd/C1st was less than this value even though the second network was synthesized at a concentration two times that in the standard DN gels (Table 2). Owing to the relatively low charge concentration of the biopolymers, the first network did not swell as much as in standard DN gels or St-DN synthetic gels, and a smaller amount of second network relative to the first network was incorporated. As final rupture in tension is determined by the balance of the two networks, the relatively low concentration of the second network leads to smaller fracture strains in comparison with standard DN gels and St-DN synthetic gels. Another possibility is increased chain entanglement within the second network, since it was formed at twice the concentration used in standard DN gels and St-DN synthetic gels.[11,14] The high entanglement of the second network leads to a small fracture strain. The increasing slope in the post-yielding region supports this explanation. It should be mentioned that when the second network was synthesized at low monomer concentration (1 M, 2 M), no yielding occurred and the St-DN gels showed brittle behavior, similar to SN gels (Figure S5, Supporting Information). We now present the results of simulated biomedical applications. For biomedical applications, the mechanical properties of the hydrogels must not change simply due to being placed in an in vivo environment. (However, subsequent changes, e.g. under long-term cyclic dynamic loads, are probably inevitable.) Thus, we compared the initial swelling degree and mechanical properties of the St(PGs)-DN gels in water and in physiological saline solution. No volume changes of the samples were observed in saline solution, and the tensile stress–strain curves of St(PGs)DN gels swelled in water and in saline solution overlapped, as shown in Figure 4a. These results indicate that the St-DN gels should be initially stable in physiological environments. For common polyelectrolyte-based hydrogels, the swelling of the gels is determined by the balance between the network elasticity and the ionic osmotic pressure. The latter strongly depends on the ionic strength of the surrounding buffer solution. In the DN hydrogels, however, the swelling of the gel is predominantly determined by the balance of the elasticity of the first network and the osmotic pressure of the second neutral network, owing to their contrasting structures. As the second network in these studies was neutral, the swelling of the tough DN gels is not influenced by the ionic strength of the buffer. To further investigate the potential of the St-DN gels as biomaterials for use in intimate contact with living tissues, we cultured several kinds of cells on DN gels without and with the biopolymers. Typical phase-contrast microscopy images of human coronary artery endothelial cells (HCAECs) cultured for 5 days on DN gels without and with PGs are shown in Figure 4b and c, respectively. HCAECs retained a spherical phenotype on the sample St(0)-DN gel that did not contain biopolymer (Figure 4b). This indicates that the neutral, bioinert PDMAAm/PDMAAm gel offers no specific biochemical ability to promote proliferation of the HCAECs. This is in agreement with the previous results that cells showed low adhesion to PDMAAm SN gels.[16] In contrast, on the St(PGs)-DN3 gels, cells adhered and assumed a spindle shape, which is the specific phenotype of endothelial cells (Figure 4c). In addition, the number of cells per unit area also increased on the St(PGs)DN3 gel in comparison with the St(0)-DN gel. These results
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indicate that PGs in the DN gel enhanced the adhesion and may promote proliferation of HCAECs cells to some extent. However, rabbit chondrocytes, human articular chondrocytes, human skin fibroblasts (TIG-3S), and mouse hippocampal neurons all showed weak adhesion to the DN gels with or without CS. These results demonstrate that biopolymers can improve the bioactivity of DN gels in some cases but are by no means universal in effect. The limited effects found may be attributable to the low biopolymer content (Table 2). A further study in which the biopolymer content in the DN gels is increased is required to possibly improve the bioactivity of the gels. This work demonstrates that the highly charged proteoglycans and glycosaminoglycans dramatically improve the strength and toughness of a neutral DN gel. Our findings also provide insight into how these bio-polyelectrolytes play a role as structural components in cartilage and other tissues.
Experimental Section Synthesis: The St-DN gels were synthesized by the two-step polymerization method.[14] The biopolymer used as a molecular stent was incorporated into the precursor solution of the first network using a planetary centrifugal vacuum mixer (Thinky Corp., Tokyo). Then the solutions were poured into a simple reaction cell we created, consisting of a pair of glass plates kept apart with silicone elastomer spacers 1.0 mm thick. The first network hydrogel containing biopolymer (St-SN gel) was synthesized by irradiation of the reaction cell with 365 nm UV radiation for 8 h in an argon environment. The St-SN gel was then swelled in the precursor solution of the second network for one day. After that, specimens were irradiated by UV light again for 8 h in an argon environment to perform the second step of polymerization. The St-DN gel thus obtained was immersed in pure water to reach equilibrium before further characterization. Staining of St gels: The PGs incorporated into the gels were identified by applying trivalent cationic Alcian blue staining solution. The stain selectively dyed PGs by electrostatic interaction with the sulfate group of CS in the PGs. There was no interaction with the neutral polymer of PDMAAm and thus it remained unstained. Mechanical stress–strain measurements: The nominal elastic modulus E, fracture stress σf, and the fracture strain εf were determined in tension using a commercial test machine (Tensilon RTC-1150A, Orientec Co., Tokyo).[14] The nominal stress was obtained by dividing the tensile force by the initial (unstretched) cross-sectional area of the specimen. The elongation rate was 100 mm min–1. Cell cultivation: Human coronary artery endothelial cells (HCAECs; Sanko Junyaku, Tokyo) in suspension (2.26 × 104 cells/cm2) were directly cultured on the gels without any modification. The cell plus gel specimens were kept at 37 °C in a humidified atmosphere of 5% CO2 for 5 days.[21] Experimental details and a more complete definition of hydrogel compositions can be found in the Supporting Information.
Supporting Information Supporting Information is available from the Wiley Online Library or from the author. Figure 4. a) Effect of physiological saline on the tensile stress–strain curves of St(PGs)-DN gels. b,c) Phase-contrast microscopy images of human coronary artery endothelial cells (HCAECs) cultured for 5 days on the DN gels without PGs (b) and with PGs (c). The sample code for (a) and (c) is St(PGs)-DN3 and for (b) St(0)-DN. Scale bar: 100 μm.
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Acknowledgements This research was financially supported by a Grant-in-Aid for Scientific Research (S) (No. 124225006) from the Japan Society for the Promotion
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www.MaterialsViews.com of Science (JSPS), and A-STEP (No. AS242Z01029Q) from Japan Science and Technology Agency. Received: July 22, 2013 Revised: August 23, 2013 Published online: October 22, 2013
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[10] S. Naficy, H. R. Brown, J. M. Razal, G. M. Spinks, P. G. Whitten, Aust. J. Chem. 2011, 64, 1007. [11] T. Nakajima, T. Kurokawa, S. Ahmed, W. Wu J. P. Gong, Soft Matter 2013, 9, 1955. [12] J. Y. Sun, X. Zhao, W. R. K. Illeperuma, O. Chaudhuri, K. H. Oh, D. J. Mooney, J. J. Visassak, Z. Suo, Nature 2012, 489, 133. [13] T. C. Suekama, J. Hu, T. Kurokawa, J. P. Gong, S. H. Gehrke, ACS Macro Lett. 2013, 2, 137. [14] T. Nakajima, H. Sato, Y. Zhao, S. Kawahara, J. J. Yang, T. Kurokawa, K. Sugahara, J. P. Gong, Adv. Funct. Mater. 2012, 22, 4426. [15] I. S. Isayewa, B. T. Kasibhatala, K. S. Rosenthal, J. P. Kennedy, Biomaterials 2003, 24, 3483. [16] Y. Nakayama, C. Kakei, A. Ishikawa, Y. M. Zhou, Y. Nemoto, K. Uchida, Bioconjugate Chem. 2007, 18, 2037. [17] Y. M. Chen, R. Ogawa, A. Kakugo, Y. Osada, J. P. Gong, Soft Matter 2009, 5, 1804. [18] K. Arakaki, N. Kitamura, T. Kurokawa, S. Onodera, F. Kanaya, J. P. Gong, K. Yasuda, J. Mater. Sci.: Mater. Med. 2011, 22(2), 417. [19] M. Rubinstein, R. H. Colby, A. V. Dobrynin, J.-F. Joanny, Macromolecules 1996, 29, 398. [20] Y.-H. Na, Y. Tanaka, Y. Kawauchi, H. Furukawa, T. Sumiyoshi, J. P. Gong, Y. Osada, Macromolecules 2006, 39, 4641. [21] J. J. Yang, Y. M. Chen, T. Kurokawa, J. P. Gong, S. Onodera, K. Yasuda, J. Biomed. Mater. Res.: Part A 2010, 95, 531.
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Proteoglycans and Glycosaminoglycans Improve Toughness of Biocompatible Double Network Hydrogels Yu Zhao, Tasuku Nakajima, Jing Jing Yang, Takayuki Kurokawa, Jian Liu, Jishun Lu, Shuji Mizumoto, Kazuyuki Sugahara, Nobuto Kitamura, Kazunori Yasuda, A. U. D. Daniels, and Jian Ping Gong* Hydrogels made from biopolymers, such as hyaluronan and collagen, are widely used in biomedical applications as cell scaffolds and artificial tissues.[1–3] These biopolymer hydrogels, which are either chemically or physically cross-linked, are usually very soft with nominal tensile moduli ranging from 10 to 104 Pa.[4] Thus, like conventional hydrogels made from synthetic single polymers, these biopolymer hydrogels are much less stiff and strong than load-bearing biological tissues. For example, cartilage has a nominal modulus of 0.1–1.0 MPa, a compressive fracture stress of several megapascals and a tearing energy of more than 1000 J m–2,[5] while hyaluronan gels typically show a fracture stress of only 105 Pa.[6] Development of biopolymer-based hydrogels with mechanical performance comparable to that of load-bearing soft tissues would greatly broaden the application of these materials as substitutes for tissues such as cartilage. The double-network (DN) principle provides a general approach to developing hydrogels with cartilage-like robustness but also high water content (ca. 90 wt%).[7] One family of tough DN hydrogels consists of poly(2-acrylamido-2-methylpropanesulfonic acid) (PAMPS) as the first network and poly(acrylamide) (PAAm) as the second network. They have tensile fracture stress and tearing fracture energy as high as 2 MPa and 4000 J m–2, respectively, comparable to or even superior to cartilage.[8–10] Many studies have shown that this DN concept is general and does not depend on specific chemical interactions between the two networks.[11,12] Thus, the DN concept should be applicable to developing tough hydrogels with biochemical functionality using combinations of various biopolymers and biocompatible polymers.
The challenge is how to synthesize the specific structure required by the DN hydrogels using biopolymers and biocompatible polymers. Synthetic polymer DN gels achieve toughness by having a topological structure very different from a conventional interpenetrating network (IPN). It consists of two interpenetrating networks that have contrasting physical structures and properties: a densely cross-linked and rigid (and thus relatively brittle) network (skeleton) of low concentration and a sparsely cross-linked soft and ductile network of high concentration.[7,8] Specifically, the molar mass of the second network is optimally 7–20 times that of the first network. For synthetic polymers, this contrasting structure is formed by a two-step sequential polymerization process: first one synthesizes a polyelectrolyte network (e.g., PAMPS) as the rigid skeleton, and then one synthesizes within this structure a neutral network (e.g., PAAm) as the soft and ductile network. Following this approach, one strategy to synthesize biopolymer-based DN gel would be to use a biopolymer as the first network and a neutral biocompatible polymer as the second network. This is because biopolymers, such as proteoglycans or glycosaminoglycans, are highly charged polyelectrolytes. Following this approach, tough DN hydrogels using chemically modified methacrylate chondroitin sulfate as the first network, and biocompatible poly(N,N-dimethyl acrylamide) (PDMAAm) as the second network have been developed recently.[13] However, this approach requires chemical cross-linking of the biopolymer, which is a complicated process. Another disadvantage is that chemical cross-linking might cause unfavorable changes in the biochemical properties of the biopolymers.
Prof. J. P. Gong Laboratory of Soft & Wet Matter Faculty of Advanced Life Science Hokkaido University Kita-10-Nishi-8, Kita-ku, Sapporo, 060-0810, Japan E-mail: [email protected] Y. Zhao Laboratory of Soft & Wet Matter Graduate School of Life Science Hokkaido University Kita-10-Nishi-8, Kita-ku, Sapporo, 060-0810, Japan Dr. T. Nakajima, Dr. J. J. Yang, Dr. T. Kurokawa Laboratory of Soft & Wet Matter Faculty of Advanced Life Science Hokkaido University Kita-10-Nishi-8, Kita-ku, Sapporo, 060-0810, Japan
J. Liu Laboratory of Soft & Wet Matter Graduate School of Science Hokkaido University Kita-10-Nishi-8, Kita-ku, Sapporo, 060-0810, Japan J. Lu, S. Mizumoto, Prof. K. Sugahara Laboratory of Proteoglycan Signaling and Therapeutics Faculty of Advanced Life Science Hokkaido University Kita-21-Nishi-11, Kita-ku, Sapporo, 001-0021, Japan Dr. N. Kitamura, Prof. K. Yasuda Department of Sports Medicine and Joint Surgery Graduate School of Medicine Hokkaido University Kita-15-Nishi-7, Kita-ku, Sapporo, 060-0821, Japan Prof. A. U. D. Daniels Biomechanics & Calorimetry Center Basel Faculty of Medicine University of Basel c/o Biozentrum, Klingelbergstrasse 50-70, 4056 Basel, Switzerland
DOI: 10.1002/adma.201303387
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Here, we present a more general approach 2nd network twork to synthesizing biopolymer-containing tough 2nd monomer onom st ne 1 network DN hydrogels. It is based on the so-called Molecular Swelling in 2nd molecular stent method that we developed stent nt 2nd solution gelation recently.[14] It can be used to combine a chemically inert, neutral, biocompatible polymer network with any polyelectrolyte biopolymer to form a tough DN gel without any chemical St gel modification of the biopolymer. A hydrogel (as-prepared, formed from a neutral synthetic polymer St gel St-DN gel coil state) swells only modestly since the polymer seg(Swollen state) ments take on a coiled conformation in water. To use a neutral polymer as the first network Molecular stent 1st network , 2nd network of a DN gel, one needs to make the network Chondroitin sulfate proteoglycans (PG), rigid, as is the case with a polyelectrolyte. To Sodium hyaluronate (HA), Monomer do this, our strategy is to use a biopolymer Chondroitin sulfate (CS) N,N-Dimethylacrylamide (DMAAm) as a molecular stent, taking advantage of its Proteoglycan polyelectrolyte nature. As shown in Figure 1, DMAAm we add the biopolymer to the precursor monChondroitin sulfate omer solution of the first neutral network. Hyaluronan After the gel formation by polymerization, Crosslinker N, N’-methylenebis(acrylamide) (MBAA) the biopolymer trapped in the neutral network exerts extra ionic osmotic pressure on the network. As a result, the neutral gel MBAA swells far more substantially, and its polymer Proteoglycan aggregate strands become extended and rigid and consequently relatively brittle, as happens in Figure 1. Schematic illustration of how polyelectrolyte biopolymers facilitate formation of the polyelectrolyte gels. After that, we can intro- contrasting DN structure from a neutral biocompatible polymer. A hydrophilic hydrogel with duce a large amount of a second neutral a charge-neutral chain exhibits only modest swelling, with the polymer strands remaining in network to form the contrasting topological coiled conformation in water. When a polyelectrolyte biopolymer is entrapped in the structure it structure required for a tough (non-brittle) exerts an extra ionic osmotic pressure on the neutral network. As a result, the neutral network DN gel. The advantages of this method are swells, and its polymer strands (pink) are fully stretched, taking on an extended conformation that 1) we can introduce the biopolymers as is the case in polyelectrolyte gels. Thus the polyelectrolyte biopolymer acts as a molecular stent. Owing to the substantial swelling of the first network, an excess amount of the second without any chemical modification; 2) difmonomer can be incorporated afterward, and the second neutral network with high concentraferent biopolymers can be introduced into tion can be synthesized. As a result, the contrasting topological structure of a tough DN gel is the DN gel simultaneously with any desired formed, and simultaneously a biopolymer has been loaded into the tough DN gels. composition, mimicking the structure of soft tissues; 3) any neutral biocompatible polymer can be used as We turn now to the stiffness of St-SN gels. The high molethe first network of a biopolymer-stented DN gel. cular weights of biopolymers make their solutions extremely Using the above-mentioned approach, we have successfully viscous. As a result, the maximum biopolymer concentration developed a series of tough DN hydrogels from biopolymers of Cstent in the precursor solution of the hydrogels was limited chondroitin sulfate proteoglycans (PGs), one of the major comto 5 wt%. Above this concentration, it was difficult to prepare ponents of cartilage, and their glycosaminoglycan components, homogeneous solutions. Figure 2b shows the volume change chondroitin sulfate (CS), and sodium hyaluronate (HA), in qv (=Vswell/Vprep) of the PDMAAm St-SN gels relative to the combination with a neutral biocompatible polymer, poly(N,Nas-prepared state for samples containing the three biopolydimethylacrylamide) (PDMAAm).[15–18] These hydrogels, which mers. A substantial increase in qv is observed with Cstent. The we call St-DN gels, exhibit mechanical properties comparable to nominal tensile modulus E of the gels also increases with the standard all-synthetic PAMPS/PAAm DN hydrogels[9,10] but are incorporation of biopolymers (Figure 2c). The tensile stiffness in addition potentially bioactive. The molecular weight and charge density of biopolymers used Table 1. Molecular weight, charge density, and origin of biopolymers in this work are shown in Table 1. The molecular stent effect of used in this work. biopolymers is demonstrated in Figure 2a. The photos show (a, left) a highly swollen PDMAAm single-network hydrogel conBiopolymers Molecular weight Charge density Sources taining a PG stent, which we call St-SN gel, and (a, right) the [Da] [mol g–1] less-swollen PDMAAm gel without the PG stent. The gel con2.98 × 10−3 salmon nasal cartilage PGs 2.71 × 106 taining PG as a molecular stent was selectively stained by Alcian 6 2.63 × 10−3 HA (1.8–2) × 10 Streptococcus equi subsp. blue, while the PDMAAm gel without PG could not be stained zooepidemicus by Alcian blue. These results demonstrate that entrapping of 5 −3 2.42 × 10 squid cartilage CS (0.7–3) × 10 PGs leads to enhanced swelling of the neutral hydrogel.
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Figure 2. Effects of biopolymers on swelling of the first network composed of the neutral polymer PDMAAm. a) The SN gel containing PGs (St(PGs)SN gel) (left) and without PGs (right) after dyeing with Alcian blue. The blue color confirms the presence of the PGs and the increased volume of the St(PGs)-SN gel demonstrates the molecular stent effect created. The swelling degree qv (b), nominal elastic modulus in tension E (c), and relative modulus En (d) of St-SN gels with various concentrations of biopolymers Cstent. qv = Vswell/Vprep, En = qvEswell/Eprep, where Vswell (Eswell) and Vprep (Eprep) are the volume (modulus) of the swollen gel and the as-prepared gel, respectively. Cstent is the feed-in value of the biopolymer in the precursor solution of the gel. Sample formulation: St-2–2.
(e.g., nominal elastic modulus E) of a bulk hydrogel is correlated to both the density (number per unit volume) of elastically effective polymer strands ve and the energy of a single strand echain, E = 3veechain.[19] The biopolymers were physically entrapped in the gel and thus they did not make a significant contribution to the stiffness, while the ve of the PDMAAm strands is inversely proportional to the swelling ratio qv. Thus, the increase of E means that echain increases with the swelling, surpassing the decreasing effect of ve. As echain is also related to the stiffness of the strand, the stiffness of the PDMAAm strand in the swollen St-SN gels relative to the as-prepared state can be estimated from the normalized nominal tensile modulus En of the St-SN gels from the following equation (in which E = Eswell), En =
Vswell Vprep
E swell E prep
= qv
E E prep
(1)
As shown in Figure 2d, En increases with the biopolymer concentration Cstent. These results demonstrate that the stiffness of each neutral strand increases with the swelling due to the ionically driven osmotic pressure πion of the biopolymers. Since πion = cionRT, where cion is the concentration of the dissociated counter ions of the biopolymers, we further investigated the molecular stent effect of biopolymers in terms of their charge concentration (Figure S2, Supporting Information). The charge 438
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concentrations of biopolymers in the precursor solutions was less than 0.07 M, about 1/10 that of the synthetic polymer stent used in our previous work.[14] Consistent with this, the relative volume change qv and modulus change En were also about 1/10 those obtained using synthetic molecular stents.[14] Among the three biopolymers, CS showed the weakest stent effect at the same polymer weight concentration Cstent (Figure 2) or charge concentration (Figure S2). This is because of a higher leakage rate of CS from the gel. Owing to its relatively low molecular weight, about 50% of CS had leaked from the gel after one day, at which time the specimens were measured (Figure S3, Supporting Information), so the actual concentration of CS was only 50% in the St(CS)-SN gels. Owing to their much higher molecular weights, negligible leakage was observed in the St(PGs)- and St(HA)-SN gels. The above results are in agreement with previous studies and confirm again that the swelling effect induced by the molecular stent is universal and independent of the chemical species of the polyelectrolyte used. We also analyzed and optimized the St-DN gel stress–strain behavior. The DN gels containing biopolymers (St-DN gels) were synthesized by immersing the St-SN gels in the precursor solution of the second network for one day and performing the radical polymerization again. The modest swelling of the St-SN gels due to the low charge concentration of the biopolymers is
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Table 2. Sample codes, compositions, and mechanical properties of St-DN gels containing biopolymers. Sample code
Formulation Cstent [wt%]
x1–y1/x2–y2
Mechanical propertiesa)
True composition a)
b)
Stent conc. [wt%]
Polymer conc. [wt%]
Water conc. [wt%]
[mol/mol]
E [MPa]
σf [MPa]
C2nd/C1sta)
εf [%]
Wf [J m–3]
St(0)-DN
0
1–2/4–0.02
0
8.8
91.2
2.99
0.07
0.19
125
0.15
St(PGs)-DN1
1
1–2/4–0.02
0.14
10.26
89.6
6.27
0.11
1.5
174
0.85
St(PGs)-DN2
3
1–2/4–0.02
0.33
10.98
86.69
8.46
0.13
0.91
195
0.87
St(PGs)-DN3
5
1–4/4–0.02
0.82
12.18
87
6.98
0.33
0.96
473
3.7
St(HA)-DN1
1
1–2/4–0.02
0.16
10.44
89.4
5.78
0.1
1.15
197
1.16
St(HA)-DN2
2
1–2/4–0.02
0.55
10.74
88.7
8.60
0.14
1.06
168
0.8
St(HA)-DN3
3
1–2/4–0.02
0.58
11.14
88.2
10.70
0.19
0.73
173
0.75
St(CS)-DN1
1
1–2/4–0.02
0.09
10.11
89.8
4.91
0.09
1.36
185
0.78
St(CS)-DN2
3
1–2/4–0.02
0.18
10.42
89.4
8.09
0.11
1.2
181
0.86
St(CS)-DN3
5
1–2/4–0.02
0.28
10.62
89.1
8.73
0.12
1.04
174
0.85
St(CS)-DN4
8
0.9–4/4–0.02
0.47
11.1
88.43
10.07
0.3
0.71
426
2.26
a)In
x1–y1/x2–y2,x and y denote the DMAAm monomer concentration [M] and the MBAA crosslinker density [mol%] with respect to the monomer, respectively. The subscripts1 and 2 denote the first and second networks, respectively. C2nd/C1st is the molar ratio of DMAAm units of the second network to that of the first network. E, σf, εf, and Wf are nominal tensile elastic modulus, fracture stress, fracture strain, and work of extension at fracture of St DN gels, respectively; b)True weight contents of biopolymers, taking into consideration the leakage over time of biopolymers.
unfavorable to inducing an excess amount of the second network relative to the first network, as required to form the contrasting structures that give DN gels their exceptional mechanical properties. In order to overcome this problem, the DMAAm monomer concentration of the second network was doubled (4 M) in comparison with the previous study.[14] The formulation, the true composition, and the mechanical properties of the St-DN gels are shown in Table 2. These samples contained about 86–91 wt% water, and the molar ratio of DMAAm of the second network to that of the first network, C2nd/C1st, showed an increasing trend with the concentration of molecular stent, reaching values of 3–11. The composition data shown in Table 2 indicate that the contrasting DN structure was indeed formed in the St-DN gels. Single-network St-SN gels, packed with linear biopolymers, were relatively low in stiffness and strength, as shown by their tensile stress–strain curves in Figure 3a. Also, stress increased monotonically with the strain, and then the gels fractured abruptly with little or no plastic deformation (brittle behavior). The fracture stress and fracture strain were less than 0.1 MPa and 60%, respectively. In contrast, the double-network St-DN gels showed quite different stress–strain behavior (Figure 3b). The fracture stress and fracture strain increased dramatically after formation of the second network. Furthermore, the fracture stress and fracture strain of St-DN gels were remarkably high compared to those made without biopolymers (Table 2). Figure 3c shows the images of a St(PGs)-DN gel withstanding 500% elongation. Also all St-DN gels showed yielding (plastic deformation) behavior before failure, a characteristic feature of tough DN gels, while no yielding was observed for specimens not containing biopolymers. Quantitatively, the work of extension at tensile fracture, a parameter that expresses one aspect of toughness of gels, increased dramatically in the case of St-DN gels (Table 2). Studies on the toughening mechanism of DN gels have shown that eventual yielding (plastic deformation)
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under tension is due to the internal fracture of the first network. Bond breakage in this network dissipates energy that would otherwise result in crack propagation, and thus imparts toughness to the DN gels.[8,11,20] Thus, the yielding behavior of the St-DN gels corresponds to the internal fracture of the PDMAAm first network and the toughening mechanism of the St-DN gels is the same as that of a conventional PAMPS/PAAm DN gel that consists of the polyelectrolyte PAMPS as the first network and the neutral polymer PAAm as the second network. It should be emphasized that the linear biopolymers, physically entrapped in the double network, only serve as molecular stents that facilitate formation of the rigid, extended structure of the first network. They do not transfer force between the two networks during deformation and thus play no role mechanically in the enhancement of the double network gels. Without the biopolymers, the contrasting topological structures could not be formed in neutral double networks, and in such a case a DN gel behaves like a conventional IPN gel, showing little or no yield under tension, and thus elastic deformation followed by brittle fracture. First-generation St-DN biopolymer gels have some limitations. Compared to standard DN gels, and St-DN gels made using a synthetic polymer as a molecular stent (St-DN synthetic gels), the fracture strain of the St-PDMAAm/PDMAAm DN gels made using biopolymer stents was 2–3 times smaller.[11,14] Furthermore, the stress–strain curves of St-DN gels showed a short flat region after yielding, but then an upward slope after that (Figure 3b). Similar behaviors were also observed in the StPAAm/PAAm DN gels, when PAAm was used as both the first and the second network with biopolymers as molecular stents (Figure S4, Supporting Information). We consider that this behavior stems from the different composition ratio C2nd/C1st of the second network to the first network in these cases. For a standard DN gel and St-DN synthetic gel, optimized C2nd/C1st was around 20. In the case of the St-DN biopolymer
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Figure 3. Effect of biopolymers on the mechanical behavior of SN gels and DN gels composed of the neutral polymer PDMAAm. a,b) Tensile stress–strain curves of single-network gels (St-SN gels) (a) and doublenetwork gels (St-DN gels) (b). c) Photographs demonstrating how a St-DN gel sustains a large strain of 500%. The sample codes of the St-DN gels used in (b) are (䊏) St(0)-DN, (ⵧ) St(PGs)-DN3, (䉭) St(HA)-DN3, (x) St(CS)-DN4. The sample code of the sample used in (c) is St(PGs)DN3. The compositions are shown in Table 1. The St-SN gels in (a) have first-network compositions corresponding to those of the St-DN gels in (b). Note that different scales are used in (a) and (b).
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gels, C2nd/C1st was less than this value even though the second network was synthesized at a concentration two times that in the standard DN gels (Table 2). Owing to the relatively low charge concentration of the biopolymers, the first network did not swell as much as in standard DN gels or St-DN synthetic gels, and a smaller amount of second network relative to the first network was incorporated. As final rupture in tension is determined by the balance of the two networks, the relatively low concentration of the second network leads to smaller fracture strains in comparison with standard DN gels and St-DN synthetic gels. Another possibility is increased chain entanglement within the second network, since it was formed at twice the concentration used in standard DN gels and St-DN synthetic gels.[11,14] The high entanglement of the second network leads to a small fracture strain. The increasing slope in the post-yielding region supports this explanation. It should be mentioned that when the second network was synthesized at low monomer concentration (1 M, 2 M), no yielding occurred and the St-DN gels showed brittle behavior, similar to SN gels (Figure S5, Supporting Information). We now present the results of simulated biomedical applications. For biomedical applications, the mechanical properties of the hydrogels must not change simply due to being placed in an in vivo environment. (However, subsequent changes, e.g. under long-term cyclic dynamic loads, are probably inevitable.) Thus, we compared the initial swelling degree and mechanical properties of the St(PGs)-DN gels in water and in physiological saline solution. No volume changes of the samples were observed in saline solution, and the tensile stress–strain curves of St(PGs)DN gels swelled in water and in saline solution overlapped, as shown in Figure 4a. These results indicate that the St-DN gels should be initially stable in physiological environments. For common polyelectrolyte-based hydrogels, the swelling of the gels is determined by the balance between the network elasticity and the ionic osmotic pressure. The latter strongly depends on the ionic strength of the surrounding buffer solution. In the DN hydrogels, however, the swelling of the gel is predominantly determined by the balance of the elasticity of the first network and the osmotic pressure of the second neutral network, owing to their contrasting structures. As the second network in these studies was neutral, the swelling of the tough DN gels is not influenced by the ionic strength of the buffer. To further investigate the potential of the St-DN gels as biomaterials for use in intimate contact with living tissues, we cultured several kinds of cells on DN gels without and with the biopolymers. Typical phase-contrast microscopy images of human coronary artery endothelial cells (HCAECs) cultured for 5 days on DN gels without and with PGs are shown in Figure 4b and c, respectively. HCAECs retained a spherical phenotype on the sample St(0)-DN gel that did not contain biopolymer (Figure 4b). This indicates that the neutral, bioinert PDMAAm/PDMAAm gel offers no specific biochemical ability to promote proliferation of the HCAECs. This is in agreement with the previous results that cells showed low adhesion to PDMAAm SN gels.[16] In contrast, on the St(PGs)-DN3 gels, cells adhered and assumed a spindle shape, which is the specific phenotype of endothelial cells (Figure 4c). In addition, the number of cells per unit area also increased on the St(PGs)DN3 gel in comparison with the St(0)-DN gel. These results
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indicate that PGs in the DN gel enhanced the adhesion and may promote proliferation of HCAECs cells to some extent. However, rabbit chondrocytes, human articular chondrocytes, human skin fibroblasts (TIG-3S), and mouse hippocampal neurons all showed weak adhesion to the DN gels with or without CS. These results demonstrate that biopolymers can improve the bioactivity of DN gels in some cases but are by no means universal in effect. The limited effects found may be attributable to the low biopolymer content (Table 2). A further study in which the biopolymer content in the DN gels is increased is required to possibly improve the bioactivity of the gels. This work demonstrates that the highly charged proteoglycans and glycosaminoglycans dramatically improve the strength and toughness of a neutral DN gel. Our findings also provide insight into how these bio-polyelectrolytes play a role as structural components in cartilage and other tissues.
Experimental Section Synthesis: The St-DN gels were synthesized by the two-step polymerization method.[14] The biopolymer used as a molecular stent was incorporated into the precursor solution of the first network using a planetary centrifugal vacuum mixer (Thinky Corp., Tokyo). Then the solutions were poured into a simple reaction cell we created, consisting of a pair of glass plates kept apart with silicone elastomer spacers 1.0 mm thick. The first network hydrogel containing biopolymer (St-SN gel) was synthesized by irradiation of the reaction cell with 365 nm UV radiation for 8 h in an argon environment. The St-SN gel was then swelled in the precursor solution of the second network for one day. After that, specimens were irradiated by UV light again for 8 h in an argon environment to perform the second step of polymerization. The St-DN gel thus obtained was immersed in pure water to reach equilibrium before further characterization. Staining of St gels: The PGs incorporated into the gels were identified by applying trivalent cationic Alcian blue staining solution. The stain selectively dyed PGs by electrostatic interaction with the sulfate group of CS in the PGs. There was no interaction with the neutral polymer of PDMAAm and thus it remained unstained. Mechanical stress–strain measurements: The nominal elastic modulus E, fracture stress σf, and the fracture strain εf were determined in tension using a commercial test machine (Tensilon RTC-1150A, Orientec Co., Tokyo).[14] The nominal stress was obtained by dividing the tensile force by the initial (unstretched) cross-sectional area of the specimen. The elongation rate was 100 mm min–1. Cell cultivation: Human coronary artery endothelial cells (HCAECs; Sanko Junyaku, Tokyo) in suspension (2.26 × 104 cells/cm2) were directly cultured on the gels without any modification. The cell plus gel specimens were kept at 37 °C in a humidified atmosphere of 5% CO2 for 5 days.[21] Experimental details and a more complete definition of hydrogel compositions can be found in the Supporting Information.
Supporting Information Supporting Information is available from the Wiley Online Library or from the author. Figure 4. a) Effect of physiological saline on the tensile stress–strain curves of St(PGs)-DN gels. b,c) Phase-contrast microscopy images of human coronary artery endothelial cells (HCAECs) cultured for 5 days on the DN gels without PGs (b) and with PGs (c). The sample code for (a) and (c) is St(PGs)-DN3 and for (b) St(0)-DN. Scale bar: 100 μm.
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Acknowledgements This research was financially supported by a Grant-in-Aid for Scientific Research (S) (No. 124225006) from the Japan Society for the Promotion
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www.MaterialsViews.com of Science (JSPS), and A-STEP (No. AS242Z01029Q) from Japan Science and Technology Agency. Received: July 22, 2013 Revised: August 23, 2013 Published online: October 22, 2013
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