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Hyaluronan scaffolds: A balance between backbone functionalization and bioactivity
Acta Biomaterialia 6 (2010) 2407–2414
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Hyaluronan scaffolds: A balance between backbone functionalization and bioactivity Doris Eng a, Michael Caplan a, Mark Preul b, Alyssa Panitch c,* a
School of Biological and Health Systems Engineering, Arizona State University, Tempe, AZ, USA Neurosurgery Research Laboratory, Barrow Neurological Institute, St. Joseph’s Hospital and Medical Center, Phoenix, AZ, USA c Weldon School of Biomedical Engineering, Purdue University, West Lafayette, IN, USA b
a r t i c l e
i n f o
Article history: Received 15 October 2009 Received in revised form 17 December 2009 Accepted 29 December 2009 Available online 4 January 2010 Keywords: Hyaluronan Hydrogels Neurite growth Scanning electron microscopy Degradation
a b s t r a c t Development of biomaterials that provide mechanical and molecular cues for wound healing and regeneration must meet several design parameters. In addition to high biocompatibility, biomaterials should possess suitable porosity as well as the ability to be chemically tailored to control parameters including biodegradability and bioactivity. These characteristics were studied in hyaluronan (HA), a natural polymer found in the body. HA was modified with thiol cross-linking sites to form a stable hydrogel scaffold to examine effects in in vitro cortical cell growth. HA with 20% and 44% thiolation was used to make gels at 0.5%, 0.75%, 1%, and 1.25% (w/v). Results indicate that the bioactivity of the HA after functionalization, as determined by degree of substitution (HA thiolation), has a greater effect on neurite outgrowth than does gel stiffness. The lower substituted HA (20%) promoted greater neurite growth as compared to the higher substituted HA (44%). Ó 2010 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction A well-designed biomaterial for wound healing applications, including nerve repair, should encompass several characteristics such as biocompatible polymerization chemistry, multiple modification strategies, tunable mechanical properties, and high bioactivity [1]. In particular, biological materials from the extracellular matrix have been utilized for scaffold design [2–4]. Of these materials, hyaluronan (HA) has been found to be suitable for hydrogel synthesis due to its versatility and biocompatibility [5–7]. More importantly, HA is found during embryonic development, suggesting that materials composed of HA may mimic conditions favorable for tissue growth and regeneration [8–11]. Hyaluronan is a nonsulfated glycosaminoglycan found ubiquitously in the extracellular matrix of tissue. It has been shown to play a vital role in cell–matrix interactions including proliferation, migration, and signaling [12]. Furthermore, HA is enriched during embryonic development but decreases in concentration as differentiation occurs – indicating a critical role of HA in early cell migration and differentiation processes [9]. This phenomenon is seen during organ morphogenesis, regeneration, and wound healing. In the central nervous system (CNS), HA is found in high levels in areas of high cellular activity such as in myelin and adult pericellular matrix and could directly impact developing neural cells [13].
* Corresponding author. Tel.: +1 765 496 1313; fax: +1 765 496 1459. E-mail address: [email protected] (A. Panitch).
HA can be covalently cross-linked into a stable three-dimensional scaffold by functional modification of the linear molecule. The degree of cross-linking and polymer content is controlled to produce hydrogels of varying mechanical strength. Shu et al. demonstrated consistent synthesis of thiolated HA with reproducible degree of substitution depending on pH and oxygen content [14]. The thiol groups on derivatized HA can be used as cross-linking sites for hydrogel formation. HA gels have been synthesized in mechanical ranges from 90 to 4000 Pa by varying cross-linking density [15]. In this regard, hydrogels of HA can be modified to match the mechanical characteristics of most soft tissues. By controlling degree of substitution or thiolation of HA, the number of bioactive sites in the HA hydrogel can be varied. Functionalization alters the chemical makeup of the HA and high degrees of substitution may modify the ability of cell receptors to recognize their natural binding ligand within the HA molecule. Interactions between HA and cells are mainly mediated by CD44, a widely distributed surface glycoprotein, and RHAMM (receptor for hyaluronic-acid-mediated motility) [16,17]. Binding to these receptors has been correlated to patterns of cell behavior such as cell migration, proliferation, and differentiation [11,12]. Past studies with HA hydrogels have shown altered cell behavior based on hydrogel properties such as stiffness and concentration; however, no correlation was found between substitution rates and cell response [6,18]. In this study, cross-linked hyaluronan scaffolds of varying HA polymer concentration and substitution rates were tested to determine influence on neurite
1742-7061/$ - see front matter Ó 2010 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2009.12.049
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outgrowth from primary cortical neurons. This cell culture model was chosen to evaluate the design of the HA hydrogels for neural tissue engineering applications. To verify physical characteristics of the HA gels, rheology, SEM, and degradation studies were performed. Biocompatibility was verified with a hemolysis assay. 2. Materials and methods 2.1. Hyaluronan modification and hydrogel preparation Hyaluronan (HA) (MW 110 kDa, Genzyme, Cambridge, MA) was thiolated according to previous methods [14,19]. Briefly, the polysaccharide was dissolved at 10 mg ml–1 in Milli-Q water (degassed) and dithiobis(propanoic dihydrazide) (DTP) was added while stirring. The pH was lowered to 4.75 by adding 1 N HCl. Next, 1ethyl-3-[3-(dimethylamino) propyl]carbodiimide (EDCI) (Sigma Chemical Co., St. Louis, MO) was added and the solution was maintained at pH 4.75. The reaction was stopped after 1 h by adding 1 M sodium hydroxide (NaOH) (VWR Scientific, West Chester, PA), raising the pH to 7.0. Then, dithiothreitol (DTT) (Sigma Chemical Co.) was added in at least 5-fold molar excess (relative to concentration of dithiobis(propanoic dihydrazide) (DTP)) and the pH was raised to 8.5 with 1 M NaOH. Finally, the pH was lowered with 1 M hydrochloric acid (HCl) (Sigma Chemical Co.) and the solution was dialyzed against HCl solution (pH 3.5, 0.3 mM HCl) with 100 mM sodium chloride (NaCl) (Sigma Chemical Co.) for 2 weeks. Substitution of the glucuronate carboxyl groups of HA by thiol groups was determined using 1H NMR (500 MHz Inova) [19]. Two substitution rates (20% and 44%) were synthesized by changing the reaction time of HA with DTP. The thiol groups on hyaluronan were used as the cross-link sites. The hydrogel formed with the addition of poly(ethylene) glycol diacrylate (PEGDA) (MW 3400 Da, SunBio PEG-Shop, South Korea), which acted as a bifunctional electrophilic cross-linker that reacted with the thiol groups via Michael-type addition. To test the effect of the gel variants on biological activity as defined by neurite extension and enzymatic degradation in vitro, HA gels at concentrations of 0.5%, 0.75%, 1%, and 1.25% (w/v) were prepared. Stock concentrations of 1.5625% (w/v) HA were prepared in phosphate buffered saline (PBS) (VWR Scientific) by solubilizing the HA and adjusting the pH to 7.4 with 1 M NaOH. A stock of 2.8125% (w/v) PEGDA was also prepared in PBS. Both solutions were filtered with a 0.2 lm syringe filter (SuporÒ Membrane, VWR Scientific) before cell culture and rheology studies. The final PEGDA/HA–DTPH ratio (w/w) was 0.45. 2.2. Rheological characterization The viscoelastic mechanical properties of the hydrogels were measured with a Physica MCR 101 rheometer (Anton Paar, Ostfildern, Germany) using a parallel plate geometry with a 20 mm diameter and 100 lm gap. The temperature of the rheometer surface was controlled at 21 °C with a built-in Peltier system. The linear range of the viscoelastic response was first measured with a frequency sweep from 0.1 to 200 rad s–1 at 0.5% strain. The gels were prepared and allowed to equilibrate overnight before running a 10 min time sweep the next morning. The time sweep was performed at an angular frequency of 0.5 rad s–1 and 0.5% strain. To prevent evaporation of the sample, an evaporation blocking chamber was lowered over the sample. For all studies, t-tests (a = 0.05) and two-tailed ANOVA (a = 0.025) were performed to determine statistical significance. 2.3. Scanning electron microscopy The morphologies of HA hydrogels were visualized by scanning electron microscopy (SEM) [19]. SEM was performed on the FEI
XL30 Field-Emission Environmental Scanning Electron Microscope (ESEM-FEG) at Arizona State University. For this study a higher substituted HA sample (50%) was used to probe differences from the lower substituted HA (20%). Fully cross-linked HA gels of 0.5%, 0.75%, 1%, and 1.25% (w/v) were flash-frozen by immersion in liquid nitrogen. Frozen gels were then fractured with a scalpel and immediately placed on a lyophilizer for freeze-drying. Gel fragments were allowed to lyophilize for at least 48 h. The HA gel was then affixed to a metal stub with carbon tape for sputter coating with gold (10 nm thick). Images were taken at 15 kV with a spot size of 2. Using a custom Matlab program, the pore area of 10 random pores from each image (n = 6) was measured by tracing around the pore. Assuming a circular pore shape, the diameter was calculated from the area. Void area calculations from the images were used to estimate porosity from the images. Using ImageJ, the void area percent was calculated by dividing the number of black pixels in the image (pore area) by the total number of pixels in the image (excluding label area) [20]. Two-tailed ANOVA analyses were used to detect significant differences between the treatment groups (a = 0.025). A post hoc Student’s t-test was used to isolate differences between two groups (a = 0.05). 2.4. Hemolysis assay A hemolysis assay was performed to verify biocompatibility of the HA gels. The method was modified from Seal and Panitch [21]. Four gels of each wt.% (0.5%, 0.75%, 1%, and 1.25%) of the 20% substituted HA and 44% substituted HA were formed (100 ll) in 1 ml syringes (VWR Scientific). Gels were allowed to equilibrate overnight at 4 °C. After gels were formed, the barrel of the syringe was cut and the gels were placed into chambered wells (Nunc LabTek II, Thermo Fisher Scientific, Waltham, MA). Triton X-100 (0.1%) (Sigma Chemical Co.) was used as positive control and 100 ll of PBS was used as negative control. Using cyanmethemoglobin standards (Stanbio Laboratory, Boerne, TX), the total blood hemoglobin concentration of bovine whole blood (Innovative Research, Novi, MI) was adjusted to 10.0 mg ml–1 using PBS. The blood was further diluted 1:8 with PBS before incubation with gels and controls. Samples were incubated with 3 ml of diluted blood for 3 h at 37 °C in a humidified incubator. To ensure exposure of the sample to the blood, the plates were rocked at low speed. Afterward, the blood was carefully removed and placed in 15 ml centrifuge tubes for centrifugation at 750g for 15 min. The supernatant was removed for analysis. To determine hemoglobin released, 0.5 ml of the supernatant was mixed with 0.5 ml of cyanmethemoglobin reagent (0.75 mM potassium cyanide and 0.6 mM potassium ferricyanide) and allowed to react for 5 min. The absorbance of the resulting solution was measured with the Omega Fluostar microplate reader (BMG LABTECH GmbH, Offenburg, Germany) at a wavelength of 540 nm. Percent hemolysis was determined against the positive control. 2.5. Degradation assay The stability of the HA gels was measured with a modified carbazole assay according to previous methods [7,22,23]. Twenty percent substituted HA and 44% substituted HA was used for this study. Each type of HA was prepared at concentrations of 0.5%, 0.75%, 1%, and 1.25% (w/v). One hundred microliter gels of each concentration were placed in chamber slides (n = 4) (Nunc LabTekII, VWR Scientific). For each concentration, four gels were incubated with hyaluronidase (50 units ml–1 PBS) at 37 °C. At selected time points 750 ll of solution was removed and replaced with equal volume of fresh enzyme solution. Time points studied were 0, 2, 4, 6, 8, and 10 h. Amount of HA degradation was measured by the release of glucuronic acid into the supernatant with the car-
D. Eng et al. / Acta Biomaterialia 6 (2010) 2407–2414
bazole assay in a 96-well plate format (VWR Scientific) [22]. The absorbance of the reaction solution was measured with the Omega Fluostar microplate reader at wavelength of 550 nm. Percent degradation was calculated across all time points for the HA gels. A repeated measures ANOVA was used to determine significant differences in percent degradation between the HA gels (a = 0.05). A post hoc Tukey’s test was used to determine differences between the groups. 2.6. Primary cortical cell culture The fragility of the HA gels and neurons required performing cell culture within a silicone microwell (Sigma Chemical Co.). As immunohistochemistry processing requires many solution changes; the culture samples were protected by confining the addition of liquid over the silicone support. Each microwell support is a silicone mold with two precut wells with a diameter of 6.6 mm (approximately the size of one well in 96-well plate). Three of the well supports were placed inside a single four-chambered slide, leaving one chamber for poly-D-lysine (Sigma–Aldrich) control culture. The control chamber was coated with 500 ll of poly–1 D-lysine (50 lg ml ) overnight at 37 °C. The viability of the cortical tissue was verified by imaging the control chambers. One fourchambered slide was used for each experimental HA group (n = 6). Plates with HA gels were prepared the day before cell culture. PEGDA was dissolved in PBS and HA–DTPH was dissolved in supplemented neurobasal media (0.5 mM glutamine and 2% B27) (Invitrogen, Carlsbad, CA). Both solutions were filtered before use. Twenty five microliters of each HA gel solution (0.5%, 0.75%, 1%, and 1.25%) was added to the microwells and allowed to form at room temperature for 2 h. Afterward, 100 ll of supplemented neurobasal media was added on top of the gels to prevent evaporation. Prepared plates with HA gels were then stored in a humidified incubator at 37 °C before cell culture the following day. Rat cortex tissue (embryonic day 18) was purchased from Brain Bits (Springfield, IL) and prepared according to recommended protocols [24,25]. To isolate neurons, the cortical tissue was transferred using a wide-bore pipette to 1 ml Hibernate E media solution (Brain Bits) containing 2 mg papain (Worthington Biochemical, Lakewood, NJ). The tissue was incubated with the papain solution at 37 °C for 30 min, and then transferred to a sterile tube containing 2% (v/v) B27/Hibernate media (Invitrogen, Carlsbad, CA). After the solution reached room temperature (5 min), the tissue was triturated 10 times over 30 s through a sterile pipette tip. The cell suspension was filtered through a 40 lm nylon filter (BD Biosciences, San Jose, CA) and collected. The filtered suspension was centrifuged at 1100 rpm for 1 min. Afterward, the supernatant was discarded and the cell pellet resuspended in 3 ml 2% B27/neurobasal media with 0.5 mM glutamine. The viability and density of the cell suspension was tested by mixing 20 ll of Trypan Blue (0.4%) (Sigma Chemical Co.) with 20 ll of the cell suspension. Cell density was counted using a hemocytometer (VWR Scientific). The final density of the cells was diluted to 2 105 cells ml–1. A plain poly-D-lysine (50 lg ml–1) (Sigma Chemical Co.) coated glass chamber was used as a control for the health of the neurons. Thirty six microliters of the cell suspension was added to the control chambers. For the experimental hydrogel groups, 6.375 ll of the cell suspension was added on top of the HA gels. After 1 h, 1 ll of supplemented neurobasal media was added to the top of the gels. 2.7. Immunohistochemistry A thorough staining process has been developed to protect the cells and the integrity of the HA gels. The cell culture chambers were used to provide a platform for inverted microscopy through the coverslip bottom. All solution changes were performed on the
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edge of each chamber over the silicone support. After 2 days of culture, the cells were fixed with warm (37 °C) 4% (v/v) paraformaldehyde (Electron Microscopy Sciences, Hatfield, PA) in phosphate buffered saline (PBS) for 1 h at room temperature. The cells were then permeabilized with 0.1% Triton X-100 solution in PBS. Afterward, Image-iT™ FX signal enhancer (Invitrogen) was applied for 2 h. Between each of these steps, each sample was rinsed with 500 ll PBS and subject to three 20-min PBS washes. An overnight blocking step with 1% bovine serum albumin (BSA) (Sigma Chemical Co.) and 10% goat serum (Invitrogen) helped ensure against nonspecific binding. Additionally, all subsequent wash steps contained 0.1% BSA in PBS. After six washes, mouse anti-bIII-tubulin (R&D Systems, Minneapolis, MN) at 5 lg ml–1 was incubated with the cells at room temperature for 2 h and overnight at 4 °C. This was followed by overnight rinsing with 0.1% BSA in PBS and addition of Alexa488-coupled goat anti-mouse F(ab)0 2-fragment secondary antibody (2 lg ml–1, Invitrogen) for 2 h at room temperature (protected from light). After this step, three washes with 0.1% BSA in PBS were performed. To visualize the nuclei, 300 nM 40 ,6-diamidino-2-phenylindole, dihydrochloride (DAPI) (Invitrogen) was incubated with the cultures. After 5 min, the DAPI was removed and the final rinse was left on the sample during imaging. 2.8. Immunofluorescence microscopy Cells were imaged with a Leica DM IRB inverted microscope (Wetzlar, Germany) equipped with a SPOT RT3 Slider Camera (Diagnostic Instruments Inc., Sterling Heights, MI) and captured using SPOT software. A USH-102DH-100W mercury lamp (USHIO America Inc., Cypress, CA) was used as the exciting light source. Two excitation ranges were used to image the anti-bIII-tubulin fluorophore and DAPI. Images were first captured in the blue excitation range (filter set I3-excitation filter BP450–490 nm and emission filter LP515 nm) with a 20 objective. Next, images were taken in the UV excitation range (filter set A-excitation filter BP 340–380 nm and emission filter LP 425 nm) at the same magnification. The images at each excitation were overlayed using ImageJ [20]. To analyze the images, custom Matlab program was used to measure the distance of each neurite from the center of the cell body. Both average and maximum neurite length were calculated for each treatment group. Two-tailed ANOVA analyses were used to detect significant differences between the treatment groups (a = 0.025). A post hoc Student’s t-test was used to isolate differences between two groups (a = 0.05). 3. Results To compare the differences between low substituted HA and highly substituted HA, several wt.% of each HA type were investigated using rheology and SEM. In addition, biological characterization was evaluated through degradation with hyaluronidase and primary rat cortical neuron culture. 3.1. Rheological characterization Rheology was performed to examine the range of mechanical strength of the HA gel concentrations. The linear viscoelastic range was determined from frequency sweeps (0.1–200 Pa) of the HA gels; strain applied to the HA gels in this region produced G* values which were fairly constant – indicating that the samples were independent of the applied frequency. Therefore, a frequency at 0.5 rad s–1 was used to determine the average G* values from time sweeps. As expected, gels with lower percent HA content had lower values of the complex modulus, G* (Fig. 1). Standard error bars for some samples are too small to be seen on the figure. The 0.5%
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ever, the difference in the diameters were only significant in the 0.5% and 1% HA groups. Similarly, in the higher substituted HA gels, lower wt.% HA gels were more porous than the higher wt.% HA gels as determined from the void area percent calculation. 3.3. Hemolysis assay The hemolysis assay demonstrated little to no hemolytic potential of the HA hydrogels. When blood was added to the gels, there was almost no hemolysis seen (see Supplemental information). In contrast, the positive control sample with Triton X-100 produced a homogenous red sample, indicating cell membrane rupture has occurred. All supernatants removed from experimental samples were as clear and colorless as PBS. Fig. 1. Complex modulus of HA gels. Error bars are standard error (n = 3).
3.4. Degradation assay gels were the weakest with values from 6 to 13 Pa. The 1.25% gels were the strongest with values from 245 to 355 Pa. Between the two HA substitution types (20% and 44%), the HA gels at each concentration were statistically different (a = 0.05). 3.2. Scanning electron microscopy Images of the gel samples revealed a porous structure with interconnected channels. In the 20% substituted HA group (Fig. 2), the pores had a more circular appearance with large, regular pores. The pattern is seen in all wt.% gels. The diameters were also consistent and ranged from 15 to 22 lm (Fig. 4). Void area percent was calculated to estimate porosity between the HA gels. Between the lower wt.% gels (0.5%, 0.75%, and 1%), there was no significant difference in void area. The void area calculation was lowest in the higher wt.% hydrogel, 1.25%, which indicated lower porosity in this gel type. Images of the 50% substituted HA gels (Fig. 3) appear to show smaller pores (10–15 lm) than the 20% substituted group; how-
HA gels were incubated at 37 °C in PBS with 50 units ml–1 of hyaluronidase (Hase) to monitor enzymatic degradation. Hyaluronidase randomly cleaves b-N-acetyl-hexosamine-(1 ? 4) glycosidic bonds in HA [26]. Release of uronic acid, measured with a modified carbazole assay, was indicative of degradation of the gel. Standard curves of known HA concentrations were determined with the corresponding modified HA. Standard error bars are depicted but some bars are too small to be visualized. HA in PBS without enzyme is slow to degrade (data not shown) and reached about 20% degradation after 6 weeks. This suggests that the thioester cross-links between the HA molecules are highly resistant to hydrolysis. The addition of hyaluronidase to the buffer solutions accelerates degradation as shown in Fig. 5. This demonstrates that, even with modifications to the HA, the enzyme can still recognize the glycosidic linkages for cleavage. A repeated measures ANOVA verified that the 20% substituted HA degraded more quickly than the 44% substituted HA (p < 0.05). This faster degradation rate was not unexpected. Biological recognition of the modified HA is retained; however, the biological recognition of the
Fig. 2. SEM images of HA gels synthesized with 20% substituted HA. (a) 0.5%, (b) 0.75%, (c) 1%, (d) 1.25%.
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Fig. 3. SEM images of HA gels synthesized with 50% substituted HA. (a) 0.5%, (b) 0.75%, (c) 1%, (d) 1.25%.
Fig. 4. (a) Measured pore diameter of HA gels. Light grey bars are 20% substituted HA and black bars are 50% substituted HA. (b) Void area of HA gels calculated from SEM images. The pore size was measured in Matlab from SEM images and the average diameter was determined. Error bars are standard error (n = 6). Asterisk represents significant difference between the two groups (a = 0.05).
Fig. 5. Degradation profile of (a) 20% substituted HA and (b) 44% substituted HA. Error bars are standard error (n = 4). A repeated measures ANOVA revealed significant difference between the two HA substitution rates. All percent polymer gel variants in the 20% substituted HA group degraded more quickly than there corresponding variants in the 44% group (p < 0.05).
44% modified gel may be reduced from that of the 20% modified HA as evidenced by the difference in degradation rates between the two gel types despite the presence of greater than 10 lM diameter pores.
3.5. Primary cortical cell culture After a culture time of two days, neurite length was calculated across all wt.% gels as average and maximum length from images
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such as shown in Fig. 6. Average length is the length measured across all neurites per neuron imaged within one treatment. Maximum length is the average length of the longest neurite measured per neuron. Standard error for all samples was calculated. Within the 20% substituted HA group, the 0.5% HA gel produced greater average length over the 0.75%, 1%, and 1.25% gels (Fig. 7a). In fact, all groups were significantly different. This result is repeated in the maximum length measurement (Fig. 7b). Within the 44% substituted HA group, the 0.5% HA produced greater average length over all groups. All groups were significantly different except between 0.75% and 1% gels. In the maximum length calculations, the 0.5% HA gel also produced the longest neurites. Between the two substituted HA groups, the 20% substituted HA gels had greater average length over the 44% substituted HA. In the maximum length measurement, all 20% gel types produced greater length over the 44% gel types except for the 1% gel group.
4. Discussion The design of hyaluronan hydrogels is being investigated for their effect on cell growth. As a highly biocompatible, natural material, hyaluronan is an attractive polymer suitable for implantation. Additionally, chemical modifications to the structure of HA
can specifically tailor its biochemical and mechanical properties. In this study, varying the wt.% concentration of HA resulted in mechanical changes in HA gels. The HA concentration also influenced the physical structure, evaluated with scanning electron microscopy, as well as the overall degradation profile in solution. Varying the degree of substitution of the HA altered the biological activity of the HA as determined by enzymatic degradation and neurite outgrowth. Overall, the study provided clues in the design of HA gels for stimulating cell growth. SEM was used to examine the internal morphologies of the different gel types. As expected, the covalent cross-links in the HA hydrogels created a porous internal structure. In the lower percent gels (0.5%, 0.75%, and 1%), the pores appear to form interconnected channels. For both 20% and 50% substituted HA, the gels appear to become more structured in the 0.75% and 1% gels forming hexagonal shapes. As the concentration increases, the pore area decreases. Overall, there is no clear trend in the internal morphologies and the percent substitution of the HA hydrogels. In other studies, investigations on pore size in collagen-glycosaminoglycan (CG) scaffolds have shown that number of attached cells is inversely related to increasing mean pore size [27]. The specific surface area in the scaffolds created by the range of pore sizes studied (95.9– 150.5 lm) strongly correlated with cell adhesion and viability, suggesting that the availability of scaffold binding sites determined cell attachment. Surface area measurements were not conducted
Fig. 6. Fluorescent images of neurons cultured on HA gels. Twenty percent substituted HA are shown in (a–d). (a) 0.5%, (b) 0.75%, (c) 1%, (d) 1.25%. Forty four percent substituted HA samples are shown in (e–h). (e) 0.5%, (f) 0.75%, (g) 1%, (h) 1.25%. Scale bar is 50 lm.
Fig. 7. (a) Average length of neurites cultured on 20% and 44% substituted HA. (b) Average of maximum length of neurites cultured on 20% and 44% substituted HA. Error bars are standard error (n = 6). Asterisk represents significant difference between the two groups (a = 0.05).
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in the present study; however, it can be speculated that the presence of more pores in the lower wt.% gels contributed to higher surface area and thus surface ligand density for cell attachment and subsequent neurite outgrowth. In neural applications, Woerly et al. investigated the ability of a porous poly(N-(2-hydroxypropyl) methacrylamide) (PHPMA) hydrogels to induce tissue formation within an in vivo model [28]. Cell infiltration as well as blood vessel formation was seen inside the hydrogel implant which contained pores from 10 to 30 lm. The range of pore diameters (10– 22 lm) in this study fits within this range for neural tissue applications. The biocompatibility of the thiolated HA hydrogels was investigated with a hemolysis assay. Cell membrane rupture and subsequent release of hemoglobin is one indication of low biocompatibility. The positive control with a nonionic detergent (Triton X-100) demonstrated how easily hemoglobin can be released into solution. The samples with HA gels produced a stable cell pellet after centrifugation with little to no measurable hemoglobin released. As expected, the hemolysis assay demonstrated that direct contact of the HA gels with whole blood resulted in low toxicity; this result is highly relevant to implantable materials. One additional benefit of HA is that past studies have shown that degradation of the HA matrix by hyaluronidase produces nontoxic byproducts to cells [7,29]. To determine the effects of both polymer density and degree of modification of the HA on biological activity and degradation, gels were exposed to hyaluronidase in vitro. Control samples, HA in phosphate buffered saline, did not degrade significantly in the time period examined (data not shown). As expected, the higher wt.% HA gels did not degrade as quickly as the lower wt.% gels, which is likely a result of reduced diffusion into the gels. It is possible that higher polymer concentrations slowed diffusion by acting as physical barriers. Also, the higher modified HA (44%) did not degrade as quickly as the lower substituted samples; this may be due to disappearance of degradable sites due to chemical modification of the HA. This study demonstrated that the HA hydrogels, though modified, still possess sites of enzymatic recognition and bioactivity. However, the degradation data cannot be explained by simple enzyme kinetics. Traditional Michaelis–Menton kinetics does not fit most of the degradation profiles. In particular, the data set for the 44% modified HA (0.5% w/v) appears to reach a plateau degradation region. If the degradation profile fit traditional enzymatic kinetics, the curve would approach 100% degradation with a defined Vmax and Km. Additionally, increasing wt.% seems to produce the same effect (see HA 20, 1.25%). The plateau region could indicate an inability of the enzyme to access the degradation sites due to locally high cross-link density and locally increased chemical modification in the high polymer density gels. The mechanical data revealed that with increasing HA wt.%, the strength of the gels increased. G*, the complex modulus, is reported here and represents both the storage and loss modulus in viscoelastic materials. G* is shown to increase with HA content, suggesting that higher wt.% are stiffer due to greater polymer content within the gel. The 44% substituted HA gels were significantly stiffer than the 20% substituted HA for all gel wt.% with the largest difference in the 1.25% gels. The increased polymer content allows for the formation of more cross-links, via disulfide bond formation, since there is an excess of free thiol on the HA over those thiols used for crosslinking. The trend for both gel types is similar; increasing HA concentration creates stiffer gels. The different gel wt.% can be categorized into different ranges: 0.5% range: 0.1–10 Pa, 0.75% range: 50–100 Pa, 1% range: 150–250 Pa, and 1.25% range: 250–400 Pa. A study by Rammensee et al. demonstrated that neurofilaments, the main structural component of axons, has a G* around 2 Pa at an oscillatory frequency under 1 Hz [30]. Also, rheology of hippocampal tissue revealed a G* of 130 Pa for the tissue and 100 Pa for the
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individual glial cells [31]. The HA gels described here have G* values within the ranges described previously for central nervous system cells and tissues. Cells in the CNS are largely shielded from mechanical forces and may prosper in an environment built of softer materials. Indeed, studies by Saha et al. demonstrated that softer gels (100–500 Pa) promoted neural growth as opposed to glial growth [32]. Additionally, cells have been shown to migrate faster on softer substrates [33]. Working with HA gels with strength less than 250 Pa may prove optimal for sustaining specific neural growth and subsequent extension of processes. Cortical cell culture of neurons on top of HA hydrogels revealed a strong correlation between gel wt.% and neurite length. As the gel polymer concentration increased, the average and maximum neurite length decreased (Fig. 7a and b). The trend is repeated in both 20% and 44% substituted HA gels. Additionally, neurite length inversely correlates with higher gel stiffness. The stiffest gels (1.25%) yielded the lowest average and maximum neurite length. This result correlates with previous studies on neurite growth, which demonstrated increased neurite branching and length on soft, deformable bisacrylamide substrates (230 Pa) over stiff substrates (550 Pa) [34]. Additionally, culture of neurons from dorsal root ganglia showed a similar response in agarose gels – increasing gel stiffness with polymer density decreased the rate of neurite growth [35]. Interestingly, the cell response in this study was able to differentiate gel stiffness in smaller ranges, from 10 Pa (0.5% HA 20) to 70 Pa (0.75% HA 20) – suggesting fine biomechanical sensing by cortical neurons may be possible. Comparing the measured neurite lengths between the two types of substituted HA highlights a larger difference. The average length for the 20% substituted HA range from 107 to 25 lm whereas the 44% HA lengths range from 59 to 17 lm. The range of the maximum neurite length between the 20% and 44% substituted HA is also significant (28–203 vs. 19–100 lm). In average and maximum length, the neurite lengths were almost twice as long in the 20% substituted HA over the 44% substituted HA gels for the same polymer density (0.5%). The stiffness between these two gel types were similar, 10 vs. 13 Pa. This suggests that biological activity has greater impact in determining neurite length than mechanical properties within the studied stiffness range. The influence of the degree of derivation suggests that biological activity was modulated by the number of available interaction sites on the HA hydrogel. Between the 20% and the 44% substituted HA, the lower modified HA hydrogel has at least twice as many binding sites as the higher modified HA. Assuming interaction between the HA hydrogel and cellular receptors is determined by a decasaccharide binding site, the higher modified HA may disrupt cell adhesion [36]. The additional binding sites in lower substituted HA may facilitate higher amount of neurite extension during the same growth period. 5. Conclusion In this study, we have shown the influence of cross-linked HA hydrogels on promoting neurite outgrowth. Previous studies have focused on specific characteristics of scaffold design; this study examines both mechanical and chemical cues to neuron growth. To investigate the physical characteristics of the HA hydrogels, we examined mechanical strength with rheology and demonstrated increasing gel stiffness with HA concentration. SEM images revealed a porous internal structure in the lower HA wt.% gels that will allow ingrowth and support cell access into the interior. The degradation profile showed how lower substituted HA hydrogels degraded more quickly, suggesting higher recognition for enzymatic cleavage and bioactivity. The hypothesis of higher bioactivity was supported by the increased neurite length on the lower substituted HA hydrogels. Although a significant difference was seen between the HA concentrations, at similar substrate stiffnesses, the
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number of binding sites may be a greater influence on neurite growth. With further addition of bioactive components, the HA hydrogels may be optimized to support neurite growth. Acknowledgments The authors gratefully acknowledge the use of facilities in the John M. Cowley Center for High Resolution Electron Microscopy and the LeRoy Eyring Center for Solid State Science at Arizona State University. We also thank Dr. Jeremy Brower for his invaluable help with Matlab programming. Funding for this work was provided by the Arizona Biomedical Research Commission (0017). Appendix A. Figures with essential colour discrimination Certain figures in this article, particularly Figure 6, are difficult to interpret in black and white. The full colour images can be found in the on-line version, at doi: 10.1016/j.actbio.2009.12.049. Appendix B. Supplementary data Supplementary data associated with this article can be found, in the online version, at doi:10.1016/j.actbio.2009.12.049. References [1] Ratner BD, Bryant SJ. Biomaterials: where we have been and where we are going. Annu Rev Biomed Eng 2004;6:41–75. [2] Willerth SM, Johnson PJ, Maxwell DJ, Parsons SR, Doukas ME, Sakiyama-Elbert SE. Rationally designed peptides for controlled release of nerve growth factor from fibrin matrices. J Biomed Mater Res A 2007;80A:13–23. [3] Willits RK, Skornia SL. Effect of collagen gel stiffness on neurite extension. J Biomater Sci Polym Ed 2004;15:1521–31. [4] Deister C, Aljabari S, Schmidt CE. Effects of collagen 1, fibronectin, laminin and hyaluronic acid concentration in multi-component gels on neurite extension. J Biomater Sci Polym Ed 2007;18:983–97. [5] Park YD, Tirelli N, Hubbell JA. Photopolymerized hyaluronic acid-based hydrogels and interpenetrating networks. Biomaterials 2003;24:893–900. [6] Leach JB, Bivens KA, Patrick J, Charles W, Schmidt CE. Photocrosslinked hyaluronic acid hydrogels: natural, biodegradable tissue engineering scaffolds. Biotechnol Bioeng 2003;82:578–89. [7] Shu X, Liu Y, Palumbo F, Lu Y, Prestwich G. In situ crosslinkable hyaluronan hydrogels for tissue engineering. Biomaterials 2004;25:1339–48. [8] Baier C, Baader SL, Jankowski J, Gieselmann V, Schilling K, Rauch U, et al. Hyaluronan is organized into fiber-like structures along migratory pathways in the developing mouse cerebellum. Matrix Biol 2007;26:348–58. [9] Meszar Z, Felszeghy S, Veress G, Matesz K, Szekely G, Modis L. Hyaluronan accumulates around differentiating neurons in spinal cord of chicken embryos. Brain Res Bull 2008;75:414–8. [10] Meszar Z, Matesz K, Szigethy ZM, Veress G, Szekely G, Felszeghy S, et al. Distribution of hyaluronan and hyaluronan-associated proteins in the spinal cord of chicken embryos. FEBS J 2005;272:273. [11] Toole BP. Hyaluronan in morphogenesis. J Intern Med 1997;242:35–40.
[12] Toole BP. Hyaluronan: from extracellular glue to pericellular cue. Nat Rev Cancer 2004;4:528–39. [13] Lynn BD, Li XB, Cattini PA, Turley EA, Nagy JI. Identification of sequence, protein isoforms, and distribution of the hyaluronan-binding protein RHAMM in adult and developing rat brain. J Comp Neurol 2001;439:315–30. [14] Shu X, Liu Y, Luo Y, Roberts M, Prestwich G. Disulfide cross-linked hyaluronan hydrogels. Biomacromolecules 2002;3:1304–11. [15] Ghosh K, Shu XZ, Mou R, Lombardi J, Prestwich GD, Rafailovich MH, et al. Rheological characterization of in situ cross-linkable hyaluronan hydrogels. Biomacromolecules 2005;6:2857–65. [16] Ponta H, Sherman L, Herrlich PA. CD44: from adhesion molecules to signalling regulators. Nat Rev Mol Cell Biol 2003;4:33–45. [17] Turley EA, Noble PW, Bourguignon LYW. Signaling properties of hyaluronan receptors. J Biol Chem 2002;277:4589–92. [18] Bencherif SA, Srinivasan A, Horkay F, Hollinger JO, Matyjaszewski K, Washburn NR. Influence of the degree of methacrylation on hyaluronic acid hydrogels properties. Biomaterials 2008;29:1739–49. [19] Pouyani T, Harbison GS, Prestwich GD. Novel hydrogels of hyaluronic-acid – synthesis, surface-morphology, and solid-state NMR. J Am Chem Soc 1994;116:7515–22. [20] Rasband WS. ImageJ. US National Institutes of Health, Bethesda, Maryland, USA. Available from: http://rsb.info.nih.gov/ij/, 1997–2005. [21] Seal BL, Panitch A. Physical matrices stabilized by enzymatically sensitive covalent crosslinks. Acta Biomater 2006;2:241–51. [22] Cesaretti M, Luppi E, Maccari F, Volpi N. A 96-well assay for uronic acid carbazole reaction. Carbohydr Polym 2003;54:59–61. [23] Bitter T, Muir HM, Modified A. Uronic acid carbazole reaction. Anal Biochem 1962;4:330–4. [24] Brewer GJ, Price PJ. Viable cultured neurons in ambient carbon dioxide and hibernation storage for a month. Neuroreport 1996;7:1509–12. [25] Brewer GJ, Torricelli JR, Evege EK, Price PJ. Optimized survival of hippocampalneurons In B27-supplemented neurobasal™, a new serum-free medium combination. J Neurosci Res 1993;35:567–76. [26] Menzel EJ, Farr C. Hyaluronidase and its substrate hyaluronan: biochemistry, biological activities and therapeutic uses. Cancer Lett 1998;131:3–11. [27] O’Brien FJ, Harley BA, Yannas IV, Gibson LJ. The effect of pore size on cell adhesion in collagen-GAG scaffolds. Biomaterials 2005;26:433–41. [28] Woerly S, Petrov P, Sykova E, Roitbak T, Simonova Z, Harvey AR. Neural tissue formation within porous hydrogels implanted in brain and spinal cord lesions: ultrastructural, immunohistochemical, and diffusion studies. Tissue Eng 1999;5:467–88. [29] Liu Y, Zheng Shu X, Prestwich GD. Biocompatibility and stability of disulfidecrosslinked hyaluronan films. Biomaterials 2005;26:4737–46. [30] Rammensee S, Janmey PA, Bausch AR. Mechanical and structural properties of in vitro neurofilament hydrogels. Eur Biophys J Biophys Lett 2007;36: 661–8. [31] Lu YB, Franze K, Seifert G, Steinhauser C, Kirchhoff F, Wolburg H, et al. Viscoelastic properties of individual glial cells and neurons in the CNS. Proc Natl Acad Sci USA 2006;103:17759–64. [32] Saha K, Keung AJ, Irwin EF, Li Y, Little L, Schaffer DV, et al. Substrate modulus directs neural stem cell behavior. Biophys J 2008;95:4426–38. [33] Ghosh K, Pan Z, Guan E, Ge SR, Liu YJ, Nakamura T, et al. Cell adaptation to a physiologically relevant ECM mimic with different viscoelastic properties. Biomaterials 2007;28:671–9. [34] Flanagan LA, Ju YE, Marg B, Osterfield M, Janmey PA. Neurite branching on deformable substrates. Neuroreport 2002;13:2411–5. [35] Balgude AP, Yu X, Szymanski A, Bellamkonda RV. Agarose gel stiffness determines rate of DRG neurite extension in 3D cultures. Biomaterials 2001;22:1077–84. [36] Tammi R, MacCallum D, Hascall VC, Pienimaki JP, Hyttinen M, Tammi M. CD44 on keratinocytes is displaced by hyaluronan decasaccharides and not hexasaccharides. J Biol Chem 1998;273:28878–88.
Contents lists available at ScienceDirect
Acta Biomaterialia journal homepage: www.elsevier.com/locate/actabiomat
Hyaluronan scaffolds: A balance between backbone functionalization and bioactivity Doris Eng a, Michael Caplan a, Mark Preul b, Alyssa Panitch c,* a
School of Biological and Health Systems Engineering, Arizona State University, Tempe, AZ, USA Neurosurgery Research Laboratory, Barrow Neurological Institute, St. Joseph’s Hospital and Medical Center, Phoenix, AZ, USA c Weldon School of Biomedical Engineering, Purdue University, West Lafayette, IN, USA b
a r t i c l e
i n f o
Article history: Received 15 October 2009 Received in revised form 17 December 2009 Accepted 29 December 2009 Available online 4 January 2010 Keywords: Hyaluronan Hydrogels Neurite growth Scanning electron microscopy Degradation
a b s t r a c t Development of biomaterials that provide mechanical and molecular cues for wound healing and regeneration must meet several design parameters. In addition to high biocompatibility, biomaterials should possess suitable porosity as well as the ability to be chemically tailored to control parameters including biodegradability and bioactivity. These characteristics were studied in hyaluronan (HA), a natural polymer found in the body. HA was modified with thiol cross-linking sites to form a stable hydrogel scaffold to examine effects in in vitro cortical cell growth. HA with 20% and 44% thiolation was used to make gels at 0.5%, 0.75%, 1%, and 1.25% (w/v). Results indicate that the bioactivity of the HA after functionalization, as determined by degree of substitution (HA thiolation), has a greater effect on neurite outgrowth than does gel stiffness. The lower substituted HA (20%) promoted greater neurite growth as compared to the higher substituted HA (44%). Ó 2010 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction A well-designed biomaterial for wound healing applications, including nerve repair, should encompass several characteristics such as biocompatible polymerization chemistry, multiple modification strategies, tunable mechanical properties, and high bioactivity [1]. In particular, biological materials from the extracellular matrix have been utilized for scaffold design [2–4]. Of these materials, hyaluronan (HA) has been found to be suitable for hydrogel synthesis due to its versatility and biocompatibility [5–7]. More importantly, HA is found during embryonic development, suggesting that materials composed of HA may mimic conditions favorable for tissue growth and regeneration [8–11]. Hyaluronan is a nonsulfated glycosaminoglycan found ubiquitously in the extracellular matrix of tissue. It has been shown to play a vital role in cell–matrix interactions including proliferation, migration, and signaling [12]. Furthermore, HA is enriched during embryonic development but decreases in concentration as differentiation occurs – indicating a critical role of HA in early cell migration and differentiation processes [9]. This phenomenon is seen during organ morphogenesis, regeneration, and wound healing. In the central nervous system (CNS), HA is found in high levels in areas of high cellular activity such as in myelin and adult pericellular matrix and could directly impact developing neural cells [13].
* Corresponding author. Tel.: +1 765 496 1313; fax: +1 765 496 1459. E-mail address: [email protected] (A. Panitch).
HA can be covalently cross-linked into a stable three-dimensional scaffold by functional modification of the linear molecule. The degree of cross-linking and polymer content is controlled to produce hydrogels of varying mechanical strength. Shu et al. demonstrated consistent synthesis of thiolated HA with reproducible degree of substitution depending on pH and oxygen content [14]. The thiol groups on derivatized HA can be used as cross-linking sites for hydrogel formation. HA gels have been synthesized in mechanical ranges from 90 to 4000 Pa by varying cross-linking density [15]. In this regard, hydrogels of HA can be modified to match the mechanical characteristics of most soft tissues. By controlling degree of substitution or thiolation of HA, the number of bioactive sites in the HA hydrogel can be varied. Functionalization alters the chemical makeup of the HA and high degrees of substitution may modify the ability of cell receptors to recognize their natural binding ligand within the HA molecule. Interactions between HA and cells are mainly mediated by CD44, a widely distributed surface glycoprotein, and RHAMM (receptor for hyaluronic-acid-mediated motility) [16,17]. Binding to these receptors has been correlated to patterns of cell behavior such as cell migration, proliferation, and differentiation [11,12]. Past studies with HA hydrogels have shown altered cell behavior based on hydrogel properties such as stiffness and concentration; however, no correlation was found between substitution rates and cell response [6,18]. In this study, cross-linked hyaluronan scaffolds of varying HA polymer concentration and substitution rates were tested to determine influence on neurite
1742-7061/$ - see front matter Ó 2010 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2009.12.049
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outgrowth from primary cortical neurons. This cell culture model was chosen to evaluate the design of the HA hydrogels for neural tissue engineering applications. To verify physical characteristics of the HA gels, rheology, SEM, and degradation studies were performed. Biocompatibility was verified with a hemolysis assay. 2. Materials and methods 2.1. Hyaluronan modification and hydrogel preparation Hyaluronan (HA) (MW 110 kDa, Genzyme, Cambridge, MA) was thiolated according to previous methods [14,19]. Briefly, the polysaccharide was dissolved at 10 mg ml–1 in Milli-Q water (degassed) and dithiobis(propanoic dihydrazide) (DTP) was added while stirring. The pH was lowered to 4.75 by adding 1 N HCl. Next, 1ethyl-3-[3-(dimethylamino) propyl]carbodiimide (EDCI) (Sigma Chemical Co., St. Louis, MO) was added and the solution was maintained at pH 4.75. The reaction was stopped after 1 h by adding 1 M sodium hydroxide (NaOH) (VWR Scientific, West Chester, PA), raising the pH to 7.0. Then, dithiothreitol (DTT) (Sigma Chemical Co.) was added in at least 5-fold molar excess (relative to concentration of dithiobis(propanoic dihydrazide) (DTP)) and the pH was raised to 8.5 with 1 M NaOH. Finally, the pH was lowered with 1 M hydrochloric acid (HCl) (Sigma Chemical Co.) and the solution was dialyzed against HCl solution (pH 3.5, 0.3 mM HCl) with 100 mM sodium chloride (NaCl) (Sigma Chemical Co.) for 2 weeks. Substitution of the glucuronate carboxyl groups of HA by thiol groups was determined using 1H NMR (500 MHz Inova) [19]. Two substitution rates (20% and 44%) were synthesized by changing the reaction time of HA with DTP. The thiol groups on hyaluronan were used as the cross-link sites. The hydrogel formed with the addition of poly(ethylene) glycol diacrylate (PEGDA) (MW 3400 Da, SunBio PEG-Shop, South Korea), which acted as a bifunctional electrophilic cross-linker that reacted with the thiol groups via Michael-type addition. To test the effect of the gel variants on biological activity as defined by neurite extension and enzymatic degradation in vitro, HA gels at concentrations of 0.5%, 0.75%, 1%, and 1.25% (w/v) were prepared. Stock concentrations of 1.5625% (w/v) HA were prepared in phosphate buffered saline (PBS) (VWR Scientific) by solubilizing the HA and adjusting the pH to 7.4 with 1 M NaOH. A stock of 2.8125% (w/v) PEGDA was also prepared in PBS. Both solutions were filtered with a 0.2 lm syringe filter (SuporÒ Membrane, VWR Scientific) before cell culture and rheology studies. The final PEGDA/HA–DTPH ratio (w/w) was 0.45. 2.2. Rheological characterization The viscoelastic mechanical properties of the hydrogels were measured with a Physica MCR 101 rheometer (Anton Paar, Ostfildern, Germany) using a parallel plate geometry with a 20 mm diameter and 100 lm gap. The temperature of the rheometer surface was controlled at 21 °C with a built-in Peltier system. The linear range of the viscoelastic response was first measured with a frequency sweep from 0.1 to 200 rad s–1 at 0.5% strain. The gels were prepared and allowed to equilibrate overnight before running a 10 min time sweep the next morning. The time sweep was performed at an angular frequency of 0.5 rad s–1 and 0.5% strain. To prevent evaporation of the sample, an evaporation blocking chamber was lowered over the sample. For all studies, t-tests (a = 0.05) and two-tailed ANOVA (a = 0.025) were performed to determine statistical significance. 2.3. Scanning electron microscopy The morphologies of HA hydrogels were visualized by scanning electron microscopy (SEM) [19]. SEM was performed on the FEI
XL30 Field-Emission Environmental Scanning Electron Microscope (ESEM-FEG) at Arizona State University. For this study a higher substituted HA sample (50%) was used to probe differences from the lower substituted HA (20%). Fully cross-linked HA gels of 0.5%, 0.75%, 1%, and 1.25% (w/v) were flash-frozen by immersion in liquid nitrogen. Frozen gels were then fractured with a scalpel and immediately placed on a lyophilizer for freeze-drying. Gel fragments were allowed to lyophilize for at least 48 h. The HA gel was then affixed to a metal stub with carbon tape for sputter coating with gold (10 nm thick). Images were taken at 15 kV with a spot size of 2. Using a custom Matlab program, the pore area of 10 random pores from each image (n = 6) was measured by tracing around the pore. Assuming a circular pore shape, the diameter was calculated from the area. Void area calculations from the images were used to estimate porosity from the images. Using ImageJ, the void area percent was calculated by dividing the number of black pixels in the image (pore area) by the total number of pixels in the image (excluding label area) [20]. Two-tailed ANOVA analyses were used to detect significant differences between the treatment groups (a = 0.025). A post hoc Student’s t-test was used to isolate differences between two groups (a = 0.05). 2.4. Hemolysis assay A hemolysis assay was performed to verify biocompatibility of the HA gels. The method was modified from Seal and Panitch [21]. Four gels of each wt.% (0.5%, 0.75%, 1%, and 1.25%) of the 20% substituted HA and 44% substituted HA were formed (100 ll) in 1 ml syringes (VWR Scientific). Gels were allowed to equilibrate overnight at 4 °C. After gels were formed, the barrel of the syringe was cut and the gels were placed into chambered wells (Nunc LabTek II, Thermo Fisher Scientific, Waltham, MA). Triton X-100 (0.1%) (Sigma Chemical Co.) was used as positive control and 100 ll of PBS was used as negative control. Using cyanmethemoglobin standards (Stanbio Laboratory, Boerne, TX), the total blood hemoglobin concentration of bovine whole blood (Innovative Research, Novi, MI) was adjusted to 10.0 mg ml–1 using PBS. The blood was further diluted 1:8 with PBS before incubation with gels and controls. Samples were incubated with 3 ml of diluted blood for 3 h at 37 °C in a humidified incubator. To ensure exposure of the sample to the blood, the plates were rocked at low speed. Afterward, the blood was carefully removed and placed in 15 ml centrifuge tubes for centrifugation at 750g for 15 min. The supernatant was removed for analysis. To determine hemoglobin released, 0.5 ml of the supernatant was mixed with 0.5 ml of cyanmethemoglobin reagent (0.75 mM potassium cyanide and 0.6 mM potassium ferricyanide) and allowed to react for 5 min. The absorbance of the resulting solution was measured with the Omega Fluostar microplate reader (BMG LABTECH GmbH, Offenburg, Germany) at a wavelength of 540 nm. Percent hemolysis was determined against the positive control. 2.5. Degradation assay The stability of the HA gels was measured with a modified carbazole assay according to previous methods [7,22,23]. Twenty percent substituted HA and 44% substituted HA was used for this study. Each type of HA was prepared at concentrations of 0.5%, 0.75%, 1%, and 1.25% (w/v). One hundred microliter gels of each concentration were placed in chamber slides (n = 4) (Nunc LabTekII, VWR Scientific). For each concentration, four gels were incubated with hyaluronidase (50 units ml–1 PBS) at 37 °C. At selected time points 750 ll of solution was removed and replaced with equal volume of fresh enzyme solution. Time points studied were 0, 2, 4, 6, 8, and 10 h. Amount of HA degradation was measured by the release of glucuronic acid into the supernatant with the car-
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bazole assay in a 96-well plate format (VWR Scientific) [22]. The absorbance of the reaction solution was measured with the Omega Fluostar microplate reader at wavelength of 550 nm. Percent degradation was calculated across all time points for the HA gels. A repeated measures ANOVA was used to determine significant differences in percent degradation between the HA gels (a = 0.05). A post hoc Tukey’s test was used to determine differences between the groups. 2.6. Primary cortical cell culture The fragility of the HA gels and neurons required performing cell culture within a silicone microwell (Sigma Chemical Co.). As immunohistochemistry processing requires many solution changes; the culture samples were protected by confining the addition of liquid over the silicone support. Each microwell support is a silicone mold with two precut wells with a diameter of 6.6 mm (approximately the size of one well in 96-well plate). Three of the well supports were placed inside a single four-chambered slide, leaving one chamber for poly-D-lysine (Sigma–Aldrich) control culture. The control chamber was coated with 500 ll of poly–1 D-lysine (50 lg ml ) overnight at 37 °C. The viability of the cortical tissue was verified by imaging the control chambers. One fourchambered slide was used for each experimental HA group (n = 6). Plates with HA gels were prepared the day before cell culture. PEGDA was dissolved in PBS and HA–DTPH was dissolved in supplemented neurobasal media (0.5 mM glutamine and 2% B27) (Invitrogen, Carlsbad, CA). Both solutions were filtered before use. Twenty five microliters of each HA gel solution (0.5%, 0.75%, 1%, and 1.25%) was added to the microwells and allowed to form at room temperature for 2 h. Afterward, 100 ll of supplemented neurobasal media was added on top of the gels to prevent evaporation. Prepared plates with HA gels were then stored in a humidified incubator at 37 °C before cell culture the following day. Rat cortex tissue (embryonic day 18) was purchased from Brain Bits (Springfield, IL) and prepared according to recommended protocols [24,25]. To isolate neurons, the cortical tissue was transferred using a wide-bore pipette to 1 ml Hibernate E media solution (Brain Bits) containing 2 mg papain (Worthington Biochemical, Lakewood, NJ). The tissue was incubated with the papain solution at 37 °C for 30 min, and then transferred to a sterile tube containing 2% (v/v) B27/Hibernate media (Invitrogen, Carlsbad, CA). After the solution reached room temperature (5 min), the tissue was triturated 10 times over 30 s through a sterile pipette tip. The cell suspension was filtered through a 40 lm nylon filter (BD Biosciences, San Jose, CA) and collected. The filtered suspension was centrifuged at 1100 rpm for 1 min. Afterward, the supernatant was discarded and the cell pellet resuspended in 3 ml 2% B27/neurobasal media with 0.5 mM glutamine. The viability and density of the cell suspension was tested by mixing 20 ll of Trypan Blue (0.4%) (Sigma Chemical Co.) with 20 ll of the cell suspension. Cell density was counted using a hemocytometer (VWR Scientific). The final density of the cells was diluted to 2 105 cells ml–1. A plain poly-D-lysine (50 lg ml–1) (Sigma Chemical Co.) coated glass chamber was used as a control for the health of the neurons. Thirty six microliters of the cell suspension was added to the control chambers. For the experimental hydrogel groups, 6.375 ll of the cell suspension was added on top of the HA gels. After 1 h, 1 ll of supplemented neurobasal media was added to the top of the gels. 2.7. Immunohistochemistry A thorough staining process has been developed to protect the cells and the integrity of the HA gels. The cell culture chambers were used to provide a platform for inverted microscopy through the coverslip bottom. All solution changes were performed on the
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edge of each chamber over the silicone support. After 2 days of culture, the cells were fixed with warm (37 °C) 4% (v/v) paraformaldehyde (Electron Microscopy Sciences, Hatfield, PA) in phosphate buffered saline (PBS) for 1 h at room temperature. The cells were then permeabilized with 0.1% Triton X-100 solution in PBS. Afterward, Image-iT™ FX signal enhancer (Invitrogen) was applied for 2 h. Between each of these steps, each sample was rinsed with 500 ll PBS and subject to three 20-min PBS washes. An overnight blocking step with 1% bovine serum albumin (BSA) (Sigma Chemical Co.) and 10% goat serum (Invitrogen) helped ensure against nonspecific binding. Additionally, all subsequent wash steps contained 0.1% BSA in PBS. After six washes, mouse anti-bIII-tubulin (R&D Systems, Minneapolis, MN) at 5 lg ml–1 was incubated with the cells at room temperature for 2 h and overnight at 4 °C. This was followed by overnight rinsing with 0.1% BSA in PBS and addition of Alexa488-coupled goat anti-mouse F(ab)0 2-fragment secondary antibody (2 lg ml–1, Invitrogen) for 2 h at room temperature (protected from light). After this step, three washes with 0.1% BSA in PBS were performed. To visualize the nuclei, 300 nM 40 ,6-diamidino-2-phenylindole, dihydrochloride (DAPI) (Invitrogen) was incubated with the cultures. After 5 min, the DAPI was removed and the final rinse was left on the sample during imaging. 2.8. Immunofluorescence microscopy Cells were imaged with a Leica DM IRB inverted microscope (Wetzlar, Germany) equipped with a SPOT RT3 Slider Camera (Diagnostic Instruments Inc., Sterling Heights, MI) and captured using SPOT software. A USH-102DH-100W mercury lamp (USHIO America Inc., Cypress, CA) was used as the exciting light source. Two excitation ranges were used to image the anti-bIII-tubulin fluorophore and DAPI. Images were first captured in the blue excitation range (filter set I3-excitation filter BP450–490 nm and emission filter LP515 nm) with a 20 objective. Next, images were taken in the UV excitation range (filter set A-excitation filter BP 340–380 nm and emission filter LP 425 nm) at the same magnification. The images at each excitation were overlayed using ImageJ [20]. To analyze the images, custom Matlab program was used to measure the distance of each neurite from the center of the cell body. Both average and maximum neurite length were calculated for each treatment group. Two-tailed ANOVA analyses were used to detect significant differences between the treatment groups (a = 0.025). A post hoc Student’s t-test was used to isolate differences between two groups (a = 0.05). 3. Results To compare the differences between low substituted HA and highly substituted HA, several wt.% of each HA type were investigated using rheology and SEM. In addition, biological characterization was evaluated through degradation with hyaluronidase and primary rat cortical neuron culture. 3.1. Rheological characterization Rheology was performed to examine the range of mechanical strength of the HA gel concentrations. The linear viscoelastic range was determined from frequency sweeps (0.1–200 Pa) of the HA gels; strain applied to the HA gels in this region produced G* values which were fairly constant – indicating that the samples were independent of the applied frequency. Therefore, a frequency at 0.5 rad s–1 was used to determine the average G* values from time sweeps. As expected, gels with lower percent HA content had lower values of the complex modulus, G* (Fig. 1). Standard error bars for some samples are too small to be seen on the figure. The 0.5%
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ever, the difference in the diameters were only significant in the 0.5% and 1% HA groups. Similarly, in the higher substituted HA gels, lower wt.% HA gels were more porous than the higher wt.% HA gels as determined from the void area percent calculation. 3.3. Hemolysis assay The hemolysis assay demonstrated little to no hemolytic potential of the HA hydrogels. When blood was added to the gels, there was almost no hemolysis seen (see Supplemental information). In contrast, the positive control sample with Triton X-100 produced a homogenous red sample, indicating cell membrane rupture has occurred. All supernatants removed from experimental samples were as clear and colorless as PBS. Fig. 1. Complex modulus of HA gels. Error bars are standard error (n = 3).
3.4. Degradation assay gels were the weakest with values from 6 to 13 Pa. The 1.25% gels were the strongest with values from 245 to 355 Pa. Between the two HA substitution types (20% and 44%), the HA gels at each concentration were statistically different (a = 0.05). 3.2. Scanning electron microscopy Images of the gel samples revealed a porous structure with interconnected channels. In the 20% substituted HA group (Fig. 2), the pores had a more circular appearance with large, regular pores. The pattern is seen in all wt.% gels. The diameters were also consistent and ranged from 15 to 22 lm (Fig. 4). Void area percent was calculated to estimate porosity between the HA gels. Between the lower wt.% gels (0.5%, 0.75%, and 1%), there was no significant difference in void area. The void area calculation was lowest in the higher wt.% hydrogel, 1.25%, which indicated lower porosity in this gel type. Images of the 50% substituted HA gels (Fig. 3) appear to show smaller pores (10–15 lm) than the 20% substituted group; how-
HA gels were incubated at 37 °C in PBS with 50 units ml–1 of hyaluronidase (Hase) to monitor enzymatic degradation. Hyaluronidase randomly cleaves b-N-acetyl-hexosamine-(1 ? 4) glycosidic bonds in HA [26]. Release of uronic acid, measured with a modified carbazole assay, was indicative of degradation of the gel. Standard curves of known HA concentrations were determined with the corresponding modified HA. Standard error bars are depicted but some bars are too small to be visualized. HA in PBS without enzyme is slow to degrade (data not shown) and reached about 20% degradation after 6 weeks. This suggests that the thioester cross-links between the HA molecules are highly resistant to hydrolysis. The addition of hyaluronidase to the buffer solutions accelerates degradation as shown in Fig. 5. This demonstrates that, even with modifications to the HA, the enzyme can still recognize the glycosidic linkages for cleavage. A repeated measures ANOVA verified that the 20% substituted HA degraded more quickly than the 44% substituted HA (p < 0.05). This faster degradation rate was not unexpected. Biological recognition of the modified HA is retained; however, the biological recognition of the
Fig. 2. SEM images of HA gels synthesized with 20% substituted HA. (a) 0.5%, (b) 0.75%, (c) 1%, (d) 1.25%.
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Fig. 3. SEM images of HA gels synthesized with 50% substituted HA. (a) 0.5%, (b) 0.75%, (c) 1%, (d) 1.25%.
Fig. 4. (a) Measured pore diameter of HA gels. Light grey bars are 20% substituted HA and black bars are 50% substituted HA. (b) Void area of HA gels calculated from SEM images. The pore size was measured in Matlab from SEM images and the average diameter was determined. Error bars are standard error (n = 6). Asterisk represents significant difference between the two groups (a = 0.05).
Fig. 5. Degradation profile of (a) 20% substituted HA and (b) 44% substituted HA. Error bars are standard error (n = 4). A repeated measures ANOVA revealed significant difference between the two HA substitution rates. All percent polymer gel variants in the 20% substituted HA group degraded more quickly than there corresponding variants in the 44% group (p < 0.05).
44% modified gel may be reduced from that of the 20% modified HA as evidenced by the difference in degradation rates between the two gel types despite the presence of greater than 10 lM diameter pores.
3.5. Primary cortical cell culture After a culture time of two days, neurite length was calculated across all wt.% gels as average and maximum length from images
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such as shown in Fig. 6. Average length is the length measured across all neurites per neuron imaged within one treatment. Maximum length is the average length of the longest neurite measured per neuron. Standard error for all samples was calculated. Within the 20% substituted HA group, the 0.5% HA gel produced greater average length over the 0.75%, 1%, and 1.25% gels (Fig. 7a). In fact, all groups were significantly different. This result is repeated in the maximum length measurement (Fig. 7b). Within the 44% substituted HA group, the 0.5% HA produced greater average length over all groups. All groups were significantly different except between 0.75% and 1% gels. In the maximum length calculations, the 0.5% HA gel also produced the longest neurites. Between the two substituted HA groups, the 20% substituted HA gels had greater average length over the 44% substituted HA. In the maximum length measurement, all 20% gel types produced greater length over the 44% gel types except for the 1% gel group.
4. Discussion The design of hyaluronan hydrogels is being investigated for their effect on cell growth. As a highly biocompatible, natural material, hyaluronan is an attractive polymer suitable for implantation. Additionally, chemical modifications to the structure of HA
can specifically tailor its biochemical and mechanical properties. In this study, varying the wt.% concentration of HA resulted in mechanical changes in HA gels. The HA concentration also influenced the physical structure, evaluated with scanning electron microscopy, as well as the overall degradation profile in solution. Varying the degree of substitution of the HA altered the biological activity of the HA as determined by enzymatic degradation and neurite outgrowth. Overall, the study provided clues in the design of HA gels for stimulating cell growth. SEM was used to examine the internal morphologies of the different gel types. As expected, the covalent cross-links in the HA hydrogels created a porous internal structure. In the lower percent gels (0.5%, 0.75%, and 1%), the pores appear to form interconnected channels. For both 20% and 50% substituted HA, the gels appear to become more structured in the 0.75% and 1% gels forming hexagonal shapes. As the concentration increases, the pore area decreases. Overall, there is no clear trend in the internal morphologies and the percent substitution of the HA hydrogels. In other studies, investigations on pore size in collagen-glycosaminoglycan (CG) scaffolds have shown that number of attached cells is inversely related to increasing mean pore size [27]. The specific surface area in the scaffolds created by the range of pore sizes studied (95.9– 150.5 lm) strongly correlated with cell adhesion and viability, suggesting that the availability of scaffold binding sites determined cell attachment. Surface area measurements were not conducted
Fig. 6. Fluorescent images of neurons cultured on HA gels. Twenty percent substituted HA are shown in (a–d). (a) 0.5%, (b) 0.75%, (c) 1%, (d) 1.25%. Forty four percent substituted HA samples are shown in (e–h). (e) 0.5%, (f) 0.75%, (g) 1%, (h) 1.25%. Scale bar is 50 lm.
Fig. 7. (a) Average length of neurites cultured on 20% and 44% substituted HA. (b) Average of maximum length of neurites cultured on 20% and 44% substituted HA. Error bars are standard error (n = 6). Asterisk represents significant difference between the two groups (a = 0.05).
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in the present study; however, it can be speculated that the presence of more pores in the lower wt.% gels contributed to higher surface area and thus surface ligand density for cell attachment and subsequent neurite outgrowth. In neural applications, Woerly et al. investigated the ability of a porous poly(N-(2-hydroxypropyl) methacrylamide) (PHPMA) hydrogels to induce tissue formation within an in vivo model [28]. Cell infiltration as well as blood vessel formation was seen inside the hydrogel implant which contained pores from 10 to 30 lm. The range of pore diameters (10– 22 lm) in this study fits within this range for neural tissue applications. The biocompatibility of the thiolated HA hydrogels was investigated with a hemolysis assay. Cell membrane rupture and subsequent release of hemoglobin is one indication of low biocompatibility. The positive control with a nonionic detergent (Triton X-100) demonstrated how easily hemoglobin can be released into solution. The samples with HA gels produced a stable cell pellet after centrifugation with little to no measurable hemoglobin released. As expected, the hemolysis assay demonstrated that direct contact of the HA gels with whole blood resulted in low toxicity; this result is highly relevant to implantable materials. One additional benefit of HA is that past studies have shown that degradation of the HA matrix by hyaluronidase produces nontoxic byproducts to cells [7,29]. To determine the effects of both polymer density and degree of modification of the HA on biological activity and degradation, gels were exposed to hyaluronidase in vitro. Control samples, HA in phosphate buffered saline, did not degrade significantly in the time period examined (data not shown). As expected, the higher wt.% HA gels did not degrade as quickly as the lower wt.% gels, which is likely a result of reduced diffusion into the gels. It is possible that higher polymer concentrations slowed diffusion by acting as physical barriers. Also, the higher modified HA (44%) did not degrade as quickly as the lower substituted samples; this may be due to disappearance of degradable sites due to chemical modification of the HA. This study demonstrated that the HA hydrogels, though modified, still possess sites of enzymatic recognition and bioactivity. However, the degradation data cannot be explained by simple enzyme kinetics. Traditional Michaelis–Menton kinetics does not fit most of the degradation profiles. In particular, the data set for the 44% modified HA (0.5% w/v) appears to reach a plateau degradation region. If the degradation profile fit traditional enzymatic kinetics, the curve would approach 100% degradation with a defined Vmax and Km. Additionally, increasing wt.% seems to produce the same effect (see HA 20, 1.25%). The plateau region could indicate an inability of the enzyme to access the degradation sites due to locally high cross-link density and locally increased chemical modification in the high polymer density gels. The mechanical data revealed that with increasing HA wt.%, the strength of the gels increased. G*, the complex modulus, is reported here and represents both the storage and loss modulus in viscoelastic materials. G* is shown to increase with HA content, suggesting that higher wt.% are stiffer due to greater polymer content within the gel. The 44% substituted HA gels were significantly stiffer than the 20% substituted HA for all gel wt.% with the largest difference in the 1.25% gels. The increased polymer content allows for the formation of more cross-links, via disulfide bond formation, since there is an excess of free thiol on the HA over those thiols used for crosslinking. The trend for both gel types is similar; increasing HA concentration creates stiffer gels. The different gel wt.% can be categorized into different ranges: 0.5% range: 0.1–10 Pa, 0.75% range: 50–100 Pa, 1% range: 150–250 Pa, and 1.25% range: 250–400 Pa. A study by Rammensee et al. demonstrated that neurofilaments, the main structural component of axons, has a G* around 2 Pa at an oscillatory frequency under 1 Hz [30]. Also, rheology of hippocampal tissue revealed a G* of 130 Pa for the tissue and 100 Pa for the
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individual glial cells [31]. The HA gels described here have G* values within the ranges described previously for central nervous system cells and tissues. Cells in the CNS are largely shielded from mechanical forces and may prosper in an environment built of softer materials. Indeed, studies by Saha et al. demonstrated that softer gels (100–500 Pa) promoted neural growth as opposed to glial growth [32]. Additionally, cells have been shown to migrate faster on softer substrates [33]. Working with HA gels with strength less than 250 Pa may prove optimal for sustaining specific neural growth and subsequent extension of processes. Cortical cell culture of neurons on top of HA hydrogels revealed a strong correlation between gel wt.% and neurite length. As the gel polymer concentration increased, the average and maximum neurite length decreased (Fig. 7a and b). The trend is repeated in both 20% and 44% substituted HA gels. Additionally, neurite length inversely correlates with higher gel stiffness. The stiffest gels (1.25%) yielded the lowest average and maximum neurite length. This result correlates with previous studies on neurite growth, which demonstrated increased neurite branching and length on soft, deformable bisacrylamide substrates (230 Pa) over stiff substrates (550 Pa) [34]. Additionally, culture of neurons from dorsal root ganglia showed a similar response in agarose gels – increasing gel stiffness with polymer density decreased the rate of neurite growth [35]. Interestingly, the cell response in this study was able to differentiate gel stiffness in smaller ranges, from 10 Pa (0.5% HA 20) to 70 Pa (0.75% HA 20) – suggesting fine biomechanical sensing by cortical neurons may be possible. Comparing the measured neurite lengths between the two types of substituted HA highlights a larger difference. The average length for the 20% substituted HA range from 107 to 25 lm whereas the 44% HA lengths range from 59 to 17 lm. The range of the maximum neurite length between the 20% and 44% substituted HA is also significant (28–203 vs. 19–100 lm). In average and maximum length, the neurite lengths were almost twice as long in the 20% substituted HA over the 44% substituted HA gels for the same polymer density (0.5%). The stiffness between these two gel types were similar, 10 vs. 13 Pa. This suggests that biological activity has greater impact in determining neurite length than mechanical properties within the studied stiffness range. The influence of the degree of derivation suggests that biological activity was modulated by the number of available interaction sites on the HA hydrogel. Between the 20% and the 44% substituted HA, the lower modified HA hydrogel has at least twice as many binding sites as the higher modified HA. Assuming interaction between the HA hydrogel and cellular receptors is determined by a decasaccharide binding site, the higher modified HA may disrupt cell adhesion [36]. The additional binding sites in lower substituted HA may facilitate higher amount of neurite extension during the same growth period. 5. Conclusion In this study, we have shown the influence of cross-linked HA hydrogels on promoting neurite outgrowth. Previous studies have focused on specific characteristics of scaffold design; this study examines both mechanical and chemical cues to neuron growth. To investigate the physical characteristics of the HA hydrogels, we examined mechanical strength with rheology and demonstrated increasing gel stiffness with HA concentration. SEM images revealed a porous internal structure in the lower HA wt.% gels that will allow ingrowth and support cell access into the interior. The degradation profile showed how lower substituted HA hydrogels degraded more quickly, suggesting higher recognition for enzymatic cleavage and bioactivity. The hypothesis of higher bioactivity was supported by the increased neurite length on the lower substituted HA hydrogels. Although a significant difference was seen between the HA concentrations, at similar substrate stiffnesses, the
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number of binding sites may be a greater influence on neurite growth. With further addition of bioactive components, the HA hydrogels may be optimized to support neurite growth. Acknowledgments The authors gratefully acknowledge the use of facilities in the John M. Cowley Center for High Resolution Electron Microscopy and the LeRoy Eyring Center for Solid State Science at Arizona State University. We also thank Dr. Jeremy Brower for his invaluable help with Matlab programming. Funding for this work was provided by the Arizona Biomedical Research Commission (0017). Appendix A. Figures with essential colour discrimination Certain figures in this article, particularly Figure 6, are difficult to interpret in black and white. The full colour images can be found in the on-line version, at doi: 10.1016/j.actbio.2009.12.049. Appendix B. Supplementary data Supplementary data associated with this article can be found, in the online version, at doi:10.1016/j.actbio.2009.12.049. References [1] Ratner BD, Bryant SJ. Biomaterials: where we have been and where we are going. Annu Rev Biomed Eng 2004;6:41–75. [2] Willerth SM, Johnson PJ, Maxwell DJ, Parsons SR, Doukas ME, Sakiyama-Elbert SE. Rationally designed peptides for controlled release of nerve growth factor from fibrin matrices. J Biomed Mater Res A 2007;80A:13–23. [3] Willits RK, Skornia SL. Effect of collagen gel stiffness on neurite extension. J Biomater Sci Polym Ed 2004;15:1521–31. [4] Deister C, Aljabari S, Schmidt CE. Effects of collagen 1, fibronectin, laminin and hyaluronic acid concentration in multi-component gels on neurite extension. J Biomater Sci Polym Ed 2007;18:983–97. [5] Park YD, Tirelli N, Hubbell JA. Photopolymerized hyaluronic acid-based hydrogels and interpenetrating networks. Biomaterials 2003;24:893–900. [6] Leach JB, Bivens KA, Patrick J, Charles W, Schmidt CE. Photocrosslinked hyaluronic acid hydrogels: natural, biodegradable tissue engineering scaffolds. Biotechnol Bioeng 2003;82:578–89. [7] Shu X, Liu Y, Palumbo F, Lu Y, Prestwich G. In situ crosslinkable hyaluronan hydrogels for tissue engineering. Biomaterials 2004;25:1339–48. [8] Baier C, Baader SL, Jankowski J, Gieselmann V, Schilling K, Rauch U, et al. Hyaluronan is organized into fiber-like structures along migratory pathways in the developing mouse cerebellum. Matrix Biol 2007;26:348–58. [9] Meszar Z, Felszeghy S, Veress G, Matesz K, Szekely G, Modis L. Hyaluronan accumulates around differentiating neurons in spinal cord of chicken embryos. Brain Res Bull 2008;75:414–8. [10] Meszar Z, Matesz K, Szigethy ZM, Veress G, Szekely G, Felszeghy S, et al. Distribution of hyaluronan and hyaluronan-associated proteins in the spinal cord of chicken embryos. FEBS J 2005;272:273. [11] Toole BP. Hyaluronan in morphogenesis. J Intern Med 1997;242:35–40.
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