Hyaluronic Acid Based Materials for Scaffolding via Two-Photon

Jan 16, 2014 - Leach, Jennie Baier; Bivens, Kathryn A.; Patrick, Charles W., Jr.; ...... Leach, Jennie B.; Bivens, Kathryn A.; Collins, Chelsea N.; Sc...
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Hyaluronic Acid Based Materials for Scaffolding via Two-Photon Polymerization Olga Kufelt,*,† Ayman El-Tamer,† Camilla Sehring,† Sabrina Schlie-Wolter,†,‡ and Boris N. Chichkov†,‡ †

Laser Zentrum Hannover e.V., Nanotechnology Department, Hollerithallee 8, 30419 Hannover, Germany Institute of Quantum Optics, Leibniz University Hannover, Welfengarten 1, 30167 Hannover, Germany

‡

ABSTRACT: Hydrogels are able to mimic the basic three-dimensional (3D) biological, chemical, and mechanical properties of native tissues. Since hyaluronic acid (HA) is a chief component of human extracellular matrix (ECM), it represents an extremely attractive starting material for the fabrication of scaffolds for tissue engineering. Due to poor mechanical properties of hydrogels, structure fabrication of this material class remains a major challenge. Two-photon polymerization (2PP) is a promising technique for biomedical applications, which allows the fabrication of complex 3D microstructures by moving the laser focus in the volume of a photosensitive material. Chemical modification of hyaluronan allows application of the 2PP technique to this natural material and, thus, precise fabrication of 3D hydrogel constructs. To create materials with tailor-made mechanochemical properties, HA was combined and covalently cross-linked with poly(ethylene glycol) diacrylate (PEGDA) in situ. 2PP was applied for the fabrication of well elaborated 3D HA and HA−PEGDA microstructures. For enhanced biological adaption, HA was functionalized with human epidermal growth factor. micromolding techniques,11,12 electrospinning,13 cell encapsulation,14 freeze-drying,15 etc. Such methods have in common that scaffolds with a defined porosity are fabricated since porosity has emerged as crucial factor allowing cell migration into the material.1,2 However, reproducibility and the variety of scaffold geometries using these methods are limited. Unlike the above-mentioned techniques, two-photon polymerization (2PP) has the unique ability to produce high resolution 3D structures within the bulk of prepolymer.16,17 Simultaneous absorption of two infrared photons initiates a chemical reaction, which produces a small volume of polymerized material. Since this phenomenon only occurs in the area of high laser intensity, that is, at the focus of a femtosecond laser, this reaction is highly localized and can occur under the material surface. A structure can be produced by scanning this focus in 3D to obtain the desired geometry with submicrometer precision. Besides the extremely high precision of this technique, no physical forces act upon the structure during fabrication unlike the droplet impacts of inkjet printing, LIFT, and MAPLE-DW or the mechanical application of a new prepolymer surface in stereolithography. In short, 2PP is the ideal technique for the fabrication of complex 3D geometries out of comparatively fragile hydrogel materials.18 In our previous publications, the fabrication of 3D scaffolds both from commercially available biologically inert materials19,20 and from biodegradable methacrylamide-modified gelatin21 using the two-photon polymerization (2PP) technique have been demonstrated. Similar to photosensitive gelatin,

1. INTRODUCTION One important task in tissue engineering and regenerative medicine is devoted to the development of scaffolds with defined geometries. Such scaffolds should mimic the natural environment of cells, enable the exchange of nutrients and metabolites, and serve as anchorage points for cell attachment in order to guide tissue formation in three dimensions (3D). The big challenges are that biocompatible scaffold materials have to be developed, which adapt perfectly to the in vivo situation but also are applicable for 3D structuring providing well-defined, variable, reproducible, and stable scaffold geometry. It is well-known that cell responses in 3D environment can differ dramatically from standard in vitro 2D culturing conditions. Therefore, the fabrication of scaffolds with tailor-made geometry and porosity has huge potential for investigations of cell behavior in 3D.1,2 During the last decades, hydrogels have attracted much attention due to their excellent physicochemical properties similar to the in vivo extracellular matrix (ECM).3 Since hyaluronic acid is a chief component of the ECM and plays a decisive role in cell proliferation,4−8 it represents a highly interesting starting material for scaffolding and biomedical applications.9,10 However, one crucial limitation in hydrogels scaffolding is their lack of stability, which complicates the production of welldefined shapes and geometries. Chemical modification of hydrogels allows the fabrication of cross-linked and, therefore, more resilient scaffold materials. The application of a locally controlled polymerization can be used to generate precisely defined structures. Regarding the variety in the material choice, numerous methods for scaffold fabrication have been described including © 2014 American Chemical Society

Received: November 20, 2013 Revised: January 9, 2014 Published: January 16, 2014 650

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described in ref 29. To introduce the vinyl groups into EGF, 1 mg of NSA was added to 500 μL of EGF dissolved in PBS (1 mg/mL) followed by incubation of the solution at 36 °C for 2 h. For purification of the conjugates, dialysis against PBS was performed for 24 h. After that, ultrafiltration was applied to concentrate the solution, which was lyophilized then. For further usage, the lyophilized product was diluted to 1 mg/mL EGF in PBS. For the fabrication of HAGM hydrogel pellets, the modified EGF was added to a solution containing 1.8 wt % Irgacure as well as 10 wt % HAGM and then polymerized under UV light for 120 min using the UV cross-linker (Bio-Link 254, Vilber Lourmat, 254 nm, 8 × 24 W) forming cylindrical samples (6 mm diameter). As control testing, hydrogel samples with unmodified EGF and equal concentrations were prepared by enclosing the growth factor physically. Therefore, the unmodified EGF was added to the polymerizable HAGM solution, followed by the preparation of cylindrical samples similar to those used for the covalently attached EGF. In both experiments, the final EGF concentration in the photopolymerizable hydrogel was 0.33 mg/mL. Prior to being tested, all samples were incubated 7 days at 4 °C in PBS, which was changed daily, to remove the photoinitiator and the unbound EGF. 2.3. Rheology. The rheology measurements were carried out using a rheometer (Physica MCR 301, Anton Paar Germany GmbH). In order to compare the viscoelastic properties of the photopolymerized HAGM bulk materials with the combined HAGM−PEG hybrids, dynamic shear oscillation experiments were applied. Further, the material properties for different polymer concentrations were verified. The viscoelastic properties of the gels were investigated using an amplitude sweep from 0.01% to 100% deformation. For this, 25 mm diameter cylindrical samples consisting of 5−15 wt % HAGM and HAGM−PEGDA, respectively, were prepared in the presence of 1.8 wt % Irgacure 2959 as photoinitiator. The samples were photopolymerized for 120 min. The operating temperature was kept constant at 37 ± 0.1 °C. These rheological measurements were performed using parallel plates of 25 mm diameter and plate-to-plate distance of 2 mm. To determine the materials’ stiffness, the complex shear modulus (G*) was calculated according Hooke’s law:30

photopolymerizable hyaluronic acid can be generated using glycidyl methacrylate-based modification. By this method HA− glycidyl methacrylate conjugates with high degrees of substitution can be obtained, which are suitable for 2PP processing and generation of porous scaffolds. Note that former studies examining photo-cross-linked HA have shown that no cell migration into the hydrogel occurred due to the lack of porosity in the material.22 In the present work, we examine chemically modified photosensitive HA materials for their suitability for microstructuring via two-photon polymerization (2PP) and fabrication of scaffolds for bone tissue engineering. Until now, the specific cell−material responses are still poorly understood; systematic studies on cell−scaffold and cell−cell interactions in a reproducible 3D hydrogel environment are important to elucidate these interactions. Therefore, biocompatibility assays were performed using osteoblasts and fibroblasts. Fibroblasts are commonly used for the standard in vitro toxicity testing referring to the ISO 109935, which is a basic requirement for the unharmful application of biomedical devices. To enhance the bioactivity of modified HA materials, epidermal growth factor (EGF) was conjugated to the material via N-succinimidyl acrylate linkage, since this biomolecule is important for cell activity including mitosis. Another enormous advantage of the 2PP technique refers to the possibility of combination of natural hydrogels with synthetic biocompatible or degradable polymers (e.g., poly(ethylene glycol) acrylates) to create new hybrid materials in situ. Because high-water content PEG hydrogel materials have already been investigated for a variety of biomedical applications, PEG was chosen as model copolymer to demonstrate this approach.23−26 Moreover, the addition of PEG provides further modulation of the material’s physicochemical properties including degradation27 and could enhance the 2PP processing time.

τ(t ) G* = γ(t )

2. MATERIALS AND METHODS All chemicals were purchased in per analysis quality and were used as received. Streptococcus equi-derived hyaluronic acid sodium salt (high molecular weight, CAS 9067-32-7), dimethylformamide (DMF), glycidyl methacrylate (GMA), and triethyl amine (TEA) were purchased from Sigma-Aldrich (Taufkirchen, Germany), as was poly(ethylene glycol) diacrylate (PEG-DA, average Mn = 6000 mol/ g). Dialysis tubes (MWCO 12 kDaA) for the purification of HA were also acquired from Sigma-Aldrich. Phosphate buffered saline (PBS), pH = 7.4 (0.0095 M PO43−, without Ca2+ and Mg2+) was obtained from Lonza (Basel, Switzerland). The photoinitiator Irgacure 2959 was purchased from Ciba AG (Basel, Switzerland). Human epidermal growth factor (EGF) (6.2 kDa globular protein) was purchased from Biochrom AG (Berlin, Germany). 2.1. Modification of Hyaluronic Acid. The unmodified hyaluronic acid was converted in a photosensitive species applying a reaction with glycidyl methacrylate (GMA) as methacrylating agent according to the protocol of S. Bencherif et al.28 Following this protocol, hyaluronic acid-glycidyl methacrylate (HAGM) conjugates were generated. Briefly, hyaluronic acid sodium salt (1 g) was dissolved in 134 mL of 0.2 M PBS and 134 mL of DMF solution (1:1), whereafter 36.64 mL of GMA and 6.43 mL of TEA were added and dissolved subsequently. After a reaction time of 10 days at room temperature, the hydrogel was precipitated twice in 500 mL of acetone. The final precipitate was dialyzed for 3 days against pure water and then lyophilized to obtain the dry hydrogel. 2.2. Epidermal Growth Factor-Modified HAGM Hydrogels. The HAGM hydrogels were further modified with human epidermal growth factor (EGF) via prior coupling with N-succinimidyl acrylate (NSA). The coupling was performed using a protocol previously

(1)

where τ is the shear stress and γ is the shear deformation. The G* moduli were calculated with according standard deviations as the average of three independently prepared and measured samples. 2.4. Swelling Measurements. The samples for investigation of swelling behavior of the HAGM and HAGM−PEGDA hydrogels, respectively, were prepared as described in section 2.2. Cylindrical hydrogel pellets with a diameter of 6 mm were stored in water for 48 h at room temperature and in PBS, alternatively. The soaked hydrogels were removed periodically. Prior to weighing, the surface water of the hydrogel cylinder was removed with a paper towel. This procedure was repeated for three samples per material, until no further weight change was detected. The swelling ratio (SWR) of the hydrogels was calculated by the method described in ref 31:

SWR(%) =

Ws − Wd × 100 Wd

(2)

where Ws is the weight of the swollen hydrogel and Wd is the hydrogel weight after freeze-drying. The results were presented as average ± SD of three identically prepared samples. 2.5. Cell Testing. Cell experiments were carried out on 6 mm diameter cylindrical pellets containing 10 wt % hydrogels and PEGDA moieties, respectively. The tests were performed on samples with the following material compositions: 10 wt % HA and 9:1 wt % HA− PEGDA. The according cylindrical samples were prepared by filling hydrogels in PDMS templates, which were enclosed between glass coverslips. The pellets were then photopolymerized for 120 min under UV light. After the samples were washed with PBS for 7 days, they were placed into a 24 well plate and sterilized under UV light for 30 651

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Figure 1. (a) Schematic illustration of the experimental setup: AOM = acousto-optical modulator, I = iris, WP = wave-plate, PM = power meter, BS = polarizing beam splitter, M = dichroic mirror, O = objective, S = sample and sample holder, X, Y, Z = positioning system. (b) For scaffolding, the laser beam was focused by an objective into the photosensitive hydrogel, which was arranged between glass slides using a PDMS spacer. min. To characterize the biocompatibility of these samples, viability and proliferation studies were performed using lactate dehydrogenase (LDH) and Alamar blue assays. Human dermal HFF-1 fibroblasts and human osteoblast-like cell line MG-63 (both DSMZ, Braunschweig, Germany) were cultivated in Dulbecco’s modified Eagles Medium F-12 (Lonza, Basel, Switzerland) supplemented with 10% fetal calf serum (FCS, Biochrom, Berlin, Germany) and antibiotics. For these studies, cells were seeded onto the samples using 2 mL of cell culture media each and incubated in a cell incubator, in which 5% CO2 atmosphere at 37 °C was maintained (Thermo Electron Cooperation, Bonn, Germany). Concerning the LDH and Alamar blue assay, additional samples of each material were cultivated under the same conditions but excluding seeded cells. These treatments referred to the blank value. For toxicity testing, the LDH assay was performed using fibroblasts and osteoblasts. This assay measures lactate dehydrogenase, a stable enzyme within the cytosol, which is only released when the dead cells undergo lysis. Therefore, high quantified values for LDH always refer to cell death. After 48 h cultivation time, 20 μL of the supernatant of each HAGM material, control, and blank was collected and treated following the online protocol of OPS diagnostics (Lebanon, NJ, USA). Right after, and additionally after 1 h incubation time at room temperature, the absorption was measured at 490 nm using a microplate reader (Infinite M2000 Pro, Tecan, Crailsheim, Germany). To quantify the amount of measured LDH, the initial absorption value, blank, and control were subtracted from the final absorption value for material. The results were given as average [intensity] ± standard error of the mean (SEM) of four independent measurements for each material. Since the results are normalized on the control and blank, intensity values bigger than 0 always refer to an increased detection of LDH, which correlates with an increase of cell death. To figure out whether the produced materials release cytotoxic fragments, eluate tests were applied using osteoblasts. For that purpose, all produced materials incubated in fresh cell culture media at 37 °C for 24 h. After that osteoblasts were seeded into a 24 well plate using the eluates for each material. Additionally, a control treatment was carried out using fresh culture media. After 48 h cultivation time, the LDH assay was performed as described. To measure the metabolic activity of living and healthy cells, the Alamar blue assay was carried out. After 48 h cultivation time of fibroblasts and osteoblasts on HAGM materials and the control, the cell culture media was replaced with 500 μL of fresh media. Then 10 μL of a 200 μg/mL resazorin stock solution was added. Since viable cells are able to convert resazorin into the fluorescent indicator resorufin during a specific time span, the measurement of fluorescent intensity indicates the metabolic activity. Similar to the LDH assay, the intensity was read out using a microplate reader (Infinite M2000 Pro, Tecan, Crailsheim, Germany) right after adding the indicator to the cell suspension and after 2 h cultivation in the cell incubator. However, the fluorescence signals were detected at 560 nm excitation and 590 nm emission. The fluorescence intensity for each material was plotted

with the blank as described. The results were given as average [intensity] ± SEM of four independent measurements for each HAGM material and the control. The effectiveness of EGF-functionalized HAGM materials was tested with fibroblasts. In contrast to the previous cell studies, culture media without FCS was used. Since FCS contains diverse growth factors, it is known that serum-free media inhibits cell growth. By counting the cell density of adherent cells on these samples under serum-free conditions after 48 and 96 h cultivation time, it could be determined whether the EGF materials are functional and stimulate proliferation in comparison to the control. The cell density was determined with a Fuchs−Rosenthal chamber. To do so, adherent cells were trypsinized using a 0.025% trypsin−EDTA−PBS solution. After a 5 min incubation time at room temperature, serum-containing medium was added, and the created cell suspension was collected and centrifuged at 500g for 10 min (Universal 320, Hettich, Düsseldorf, Germany). After that, the pellet was dissolved in medium with an appropriate volume and counted. The cell densities were normalized on the initial seeded cell density of 2.4 × 104 cells/mL and given as average in percent ± SEM of four independent measurements. 2.6. Two-Photon Polymerization: Experimental Setup. In this work, we used ytterbium femtosecond laser pulses (250 fs, 21 MHz, 520 nm) to initiate the 2PP process within the photosensitive hydrogel materials. The typical two-photon polymerization setup is shown in Figure 1a. A femtosecond laser oscillator (HIGH Q Laser, Rankweil, Austria) was used as a laser source emitting green laser radiation with an average laser output power of at least 300 mW. As a fast laser shutter, an acousto-optical modulator was used. In combination with a subsequent iris, this arrangement is required to switch the laser beam on and off. A half-wave plate in combination with a polarizing beam splitter was used to control the laser power. For scaffold fabrication, the laser beam was focused into the photosensitive material using an Epiplan 20× microscope objective (Zeiss, Oberkochen, Germany) with a numerical aperture NA = 0.4. By movement of the laser focus within the bulk of the prepolymer by means of precise positioning system (Aerotech Inc., Pittsburgh, USA), polymerization of material along the trace is obtained. 2.7. Scaffold Fabrication. For scaffold fabrication, 10 wt % hydrogels were prepared by combining Irgacure 2959 (1.8 wt %) stock solutions with the lyophilized HAGM hydrogels and mixtures containing PEGDA. The viscous liquids were enclosed between glass discs using a polydimethylsiloxane (PDMS) spacer (see Figure 1b). The following weight percent ratios were applied for 2PP processing: 10 wt % HAGM and 9:1 wt % HAGM/PEGDA. In case of a pure HAGM hydrogel, the laser parameters were adjusted at 300 μm/s for the scanning speed and 5 mW power using a 20× objective. As final step, the unpolymerized material was removed by washing the structure with distilled water. For 2PP structuring of combined hydrogels, the speed was set at 1000 μm/s and 3 mW working power. The hydrogel structures were fabricated by layer-by-layer technique producing the first layer on top 652

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Figure 2. 1H NMR spectrum of the hyaluronan−glycidyl methacrylate conjugate (HAGM) with 60% degree of substitution in D2O (olefinic protons, 6.09 and 5.66 ppm; anomeric protons, 4.46 and 4.36; carbohydrate backbone, 3.75−3.21 ppm; acetyl CH3, 1.93 ppm; methacrylate CH3, 1.86 ppm). of the lower coverslip. After each fabricated layer, the linear stage moved the sample down by 5 μm. The development of the fabricated structures was performed in distilled water until the unpolymerized material was removed. 2.8. Statistical Analysis. All quantitative data are expressed as means ± standard deviations (SD) and as standard error of the mean (SEM). Statistical analysis was performed using Student’s t-test with significant levels of p < 0.05, p < 0.01, and p < 0.001.

was used. The storage modulus (G′) and loss modulus (G″) of the cross-linked hydrogels were determined by oscillatory rheological measurements. This was performed by varying the amplitude of deformation in a particular series of trials. In short, the amplitude of deformation was changed to higher torques, while the frequency was kept constant at 1 Hz. This kind of rheological experiment is also called strain sweep. For all hydrogels, G′ > G″ was found, demonstrating the gel character of the materials (Figure 3a). The increase in the value of torque and, consequently, the degree of deformation γL led to the following results for all hydrogels (see Figure 3b). For up to 5 wt % of HAGM and HAGM−PEGDA, respectively, critical strain values in the range of 6−10% were detected. In case of the combined hydrogels with a total mass of 5 wt %, a decrease in γL was found for high amounts of PEGDA. Thus, the region of linear-viscoelastic behavior can be varied using different HAGM−PEGDA ratios. Concentrated hydrogels with a total mass of 15 wt % show γL in the range of 1%, indicating that the linear-viscoelastic region (LVR) is valid only for low deformation degrees. To evaluate the materials’ stiffness, the complex shear modulus (G*) was calculated according to section 2.3. The results for all hydrogels are shown in Figure 3c. In case of pure HAGM hydrogels, the materials’ stiffness is mainly dependent on the hydrogel concentration. The G* modulus increases from 2.6 to 12.6 kPa with increasing HAGM concentration. For the combined 5 wt % HAGM−PEGDA materials, the found G* modulus is constantly low in the range of 3 kPa independently on the amount of the incorporated PEGDA. Accordingly, the increase of the total mass to 10 wt % leads to the increased G* values of

3. RESULTS 3.1. Synthesis of the Hyaluronic Acid−Glycidyl Methacrylate Conjugates (HAGM). Photosensitive hyaluronic acid was prepared starting from hyaluronic acid sodium salt isolated from Streptococcus equi. To convert it in a photopolymerizable material, a chemical modification due to the previously described protocol28 was performed by addition of methacrylate moieties to HA backbone to generate HA− glycidyl methacrylate (HAGM) conjugates. According to this protocol HAGM yield of 88% was obtained; in our experiments the maximum degree of substitution (DS) of 60% could be achieved, verified using 1H NMR spectroscopy in deuterated water (Figure 2). The DS was calculated from the ratio of the N-acetyl peak (1.93 ppm, 3H) and the peaks of the olefinic protons (6.09 and 5.55 ppm, 2H). The successful EGF modification was verified via IR spectroscopy of the purified and lyophilized conjugate. The characteristic band for Nsubstituted amides was observed at 1738 cm−1 after the reaction with NSA, whereas bands at 979 and 876 cm−1 confirmed the introduction of vinyl groups into EGF. 3.2. Rheological Measurements. To analyze the viscoelastic properties of the materials, a rotational rheometer 653

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Figure 3. (a) Storage (G′) and loss moduli (G″) of HAGM hydrogel combined with PEGDA, respectively. (b) Degree of deformation γL for different HAGM and HAGM−PEGDA compositions. The γL values were calculated as the mean of all data points in the linear-viscoelastic region and represent the critical strain level until the sample’s network is intact. The graph shows average γL values and standard deviations of three independently and identically prepared samples. (c) Complex shear moduli (G*) for the different HAGM hydrogels and PEGDA−hybrid materials using eq 2 representing the materials’ stiffness. The graph shows average values and standard deviations of three independently and identically prepared samples. Student’s t-test analysis: p < 0.05 (∗ or +), p < 0.01 (∗∗ or ++); ∗ indicates the difference compared within each triad, and + indicates the difference compared with 5%, 10%, or 15% HAGM, respectively.

3.4. Biocompatibility of HAGM Materials. Biocompatibility of the produced materials was verified via toxicity testing, analyzing the metabolic activity and proliferation. The LDH assay revealed that all materials slightly increased the relative intensity of LDH for fibroblasts. Values of 0.09 for 5% HAGM, of 0.03 for 10% HAGM, and 0.05 for HAGM−PEG were obtained. Osteoblasts reduced the relative LDH values to −0.13 for 5% HAGM and 10% HAGM and −0.1 for HAGM−PEG, respectively. (Figure 5). Whether the produced materials degrade and release toxic fragments into the cell culture medium was determined with an eluate test. When osteoblasts were cultured in the material supernatants for 48 h, it was found that all supernatants increase the relative intensity for LDH. The biggest increase was observed for HAGM−PEG with a value of 0.2, followed by

up to 20.9 kPa indicating that the percentage of PEGDA shifts the G* moduli to larger values. In case of the 15 wt % hybrid hydrogels, the PEGDA content dominates over the material properties resulting in high G* moduli between 30.9 and 53 kPa, which is especially remarkable compared with the behavior of the 15 wt % HAGM hydrogel (12.6 kPa). 3.3. Swelling Properties. For hydrogels prepared with different HAGM−PEGDA compositions, the swelling ratio in distilled water was determined in the range of 1300% to 2700% (Figure 4). In case of 10 wt % compositions, the swelling ratio was determined to be approximately in the range of 2500%. For high amounts of HAGM−PEGDA (15 wt %), the degree of swelling is decreased to 1300%. For swelling properties in PBS, the swelling ratio in the range of 1400% was found for all materials (data not shown). 654

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Figure 4. Swelling ratios (SWR) for HAGM hydrogels and hybrid HAGM−PEGDA materials with different polymer ratios. For each group, three identically prepared samples were examined and presented as average ± SD. Student’s t-test analysis: *p < 0.05 compared with 10% HAGM; #p < 0.05 compared with 5%/5% HAGM−PEGDA; **p < 0.01 compared with 10% HAGM; ‡p < 0.01 compared with 5%/5% HAGM−PEGDA and with 2%/8% HAGM− PEGDA; ***p < 0.001 compared with 5%/5% HAGM−PEGDA.

Figure 6. Eluate testing of different HAGM materials using osteoblasts. The LDH assay was carried out after 48 h cultivation time. The obtained results are normalized on the control and presented as average ± SEM of four independent measurements. No significant difference was observed between any groups according Student’s t-test analysis.

Figure 5. Analysis of cytotoxic effects of the different HAGM materials using the LDH assay. Measurements were performed with fibroblasts and osteoblasts after 48 h cultivation time. The results are normalized on the control values and presented as average ± SEM of four independent measurements. No significant difference was observed between any groups according Student’s t-test analysis.

5% HAGM with a value of 0.1, and 10% HAGM with 0.08, respectively (Figure 6). To quantify the metabolic activity of fibroblasts and osteoblasts cultivated on different HAGM materials, the Alamar blue assay was carried out (Figure 7). On the control, fibroblasts had a relative activity of 495 and osteoblasts of 6533. On 5% HAGM and HAGM−PEG, fibroblasts slightly reduced their activity to 439 and 299, respectively; while it was increased on 10% HAGM to 1247. Osteoblasts increased their control value in comparison to the materials on 5% HAGM to 6813 but reduced it to 6176 on 10% HAGM and to 5512 on HAGM−PEG. The functionality of EGF-modified HAGM materials was analyzed with fibroblasts, quantifying their cell growth under serum-free conditions. Figure 8 shows that under control

Figure 7. Metabolic activity of (a) fibroblasts and (b) osteoblasts on HAGM materials via Alamar blue assay after 48 h cultivation time. The relative intensity is shown given as average ± SEM of four independent measurements. No significant difference was observed between any groups using Student’s t-test analysis.

conditions no cell proliferation was possible. After 96 h they reached 140%. On unmodified HAGM, the cell density was 655

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Figure 8. Cell proliferation of fibroblasts cultivated on EGF-modified HAGM under serum-free conditions. The results are normalized on the initial seeding cell density of 2.4 × 104 cells/mL and given as average in percent ± SEM of four independent measurements. Student’s t-test analysis: (*) p < 0.05 for the difference compared with the control.

even more reduced to 68%. Higher cell densities were observed when HAGM was modified with EGF. For the chemical approach, 121% fibroblasts were presented on the surface after 48 h; the physical modification resulted in a cell density of 177%. 3.5. Scaffold Fabrication Using Two-Photon Polymerization (2PP). Scaffolds with different geometries were fabricated by direct femtosecond laser writing using twophoton polymerization (2PP). According to laser radiation at around 515 nm, application of the photoinitiator Irgacure 2959 was favorable, since its absorption maximum matches the halfwavelength of the applied laser radiation. In Figure 9, several optical microscopic images are presented showing 2PP-fabricated methacrylate-modified hyaluronan structures. The images demonstrate very well that the calculated computer-aided design (CAD) models could be transferred into free-standing 3D hydrogel structures. Both ring-like and grid structures can be produced on demand with the targeted cell application. Scaffolds with predefined pore sizes between 7 and 325 μm and with dimensions of up to 2 mm were prepared (Table 1). For HAGM hydrogel scaffolding, laser power of 5 mW and 300 μm/s scanning speed were applied. After the development in distilled water including the elimination of unpolymerized material, the structures retain their shape and provide the aimed porosity. According to the fabrication procedure of the HAGM constructs, in a similar manner combined HAGM−PEG structures were generated in situ, applying the 2PP technique (Figure 10). For the combined materials, Irgacure 2959 was also used to initiate two-photon polymerization. In comparison to the 2PP processing of pure HAGM, the addition of PEGDA improved material processability and allowed increasing the scanning speed of up to 1000 μm/s, which was tested by speed power array (see Figure 10a). In this experiment, 5 × 4 scaffold structures with dimensions of 100 μm were generated by decreasing laser power horizontally from left to right, starting with 6 to 2 mW by steps of 1 mW. The scanning speed was regulated down from 2000 to 250 μm being divided by two for each new series, respectively. Well-structured samples could be

Figure 9. Microscopic images of 2PP-manufactured hyaluronic acid (HAGM) scaffold geometries: (a) one-layered ring-like structures according to the applied CAD models; (b) grid structures with different pore sizes (7, 114, and 325 μm; from left to right, light microscope 20×).

Table 1. Pore Sizes (μm) of the 2PP-Generated HAGM Structures ring structures

grid structures

46 54 84

7 114 325

Figure 10. (a) Speed power array of fabricated scaffolds consisting of combined HAGM−PEG material according to section 2.6; horizontal left to right, laser power is changed from 6 to 2 mW by steps of 1 mW; vertical top down, scanning speed is changed from 2000 to 250 μm/s by steps of 1000, 500, and 250 μm/s. (b) precisely structured HAGM−PEG grid made of 50 layers separated by 5 μm (light microscope images 20×). 656

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crucial factors in cell fate; both can influence cell behavior significantly.33 Via the formation of HAGM−PEG blends, materials with the desired elastic properties and stiffness can be produced to trigger specific osteoblast responses. On the whole, hybrid materials with a tailor-made linear-viscoelastic region can be designed for 2PP applications. 4.3. Swelling Properties of the Photo-cross-linked Materials. Different hydrogel compositions also have a great influence on the material swelling properties in water. Obviously, the 10 wt % HAGM materials show considerably higher swelling ratios than the samples consisting of 15 wt % hydrogel (Figure 4). This can be explained by the fact that cross-linked hydrogels with lower concentrations have more inner freedom to take up water than highly concentrated materials. The addition of PEGDA resulted in a slight increase from 2300% to 2700% for the swelling properties in comparison to pure HAGM materials. In PBS all materials presented a comparable swelling ratio of 1400%, which indicates that the behavior of predefined hydrogel structures is predictable in body fluids. 4.4. Biocompatibility of the Photo-cross-linked Materials. To characterize the biocompatibility for bone regeneration applications, different methods with fibroblasts and osteoblasts as a reference cell type were performed on the generated materials. Concerning the toxicity measurements, the results were cell specific. Fibroblasts slightly reacted cytotoxic to the materials, while osteoblasts did not (Figure 5). Despite the fact that the obtained results were not significant and no clear material preference could be pointed out, the presented hydrogels were shown to be suitable for the envisaged application. The slight cytotoxic effect might be related to the problematic usage of photoinitiators to start the polymerization process which was revealed in a recent study. Released into the cell culture media, these molecules may be able to cause cell death.19 To get more insight, eluate testing was carried out with osteoblasts. Figure 6 shows that all three supernatants of 5% and 10% HAGM and HAGM−PEG slightly increase the amount of dead cells. The biggest extent was observed for HAGM−PEG. However, these effects were not significant. More analyses are necessary to figure out whether the washing procedure should be lengthened or the used photoinitiators or others should be substituted and modified. Also the HA backbone functionalization degree of 60% can be important including a considerable change in the materials’ biochemistry. Studies on thiolated HA have shown that the degree of derivatization has a great impact on neurite growth, assuming more bioactive sites in HA for cell interaction in case of less modified materials.34 Comparing the metabolic activity of both cell types on the HAGM samples, no significant differences with respect to the control were obtained (Figure 7). This means that independently from the cell source all materials presented an environment that enabled high viability and metabolic activity. This result is in contradiction to the LDH toxicity assay (Figure 5), because due to the slight increase in LDH, a reduction of the metabolic activity would have been expected. Since this was not the case, we assume that all materials offer perfect conditions to be used for bone tissue engineering, as long as toxic components like the photoinitiator are washed out. 4.5. Bioactivity of the Growth Factor-Modified Gels. For successful cell applications, the HAGM hydrogels require further modification with bioactive molecules to mediate cell

obtained for laser power values of 4 and 3 mW, working at 1000 and 500 μm/s scanning speeds.

4. DISCUSSION Due to their excellent biochemical properties, hydrogels are prospective starting materials for scaffold fabrication and tissue engineering. Particularly, hyaluronic acid (HA), which is the main component of extracellular matrix (ECM), was strongly investigated during the last years. In the human body, HA is responsible for the tissue integrity being presented in various body fluids. The HA backbone provides important binding sites for cell attachment and, therefore, could be used to promote controlled cell arrangement and guided tissue formation. One of the main limiting factors for using hydrogels as supporting material is their poor mechanical properties. In this work, methacrylate-modified hyaluronic acid (HAGM) was used to fabricate well-defined cross-linked hydrogel structures as biomatrices for bone tissue engineering applications using two-photon polymerization (2PP). 4.1. Preparation of Photocurable Hyaluronic Acid. By incorporating glycidyl methacrylate groups into the HA backbone, we obtained a polymerizable hydrogel with a yield of 88% presenting a 60% DS according to 1H NMR data (Figure 2). The work of Bencherif et al. clearly shows that the amount of incorporated cross-linkable side groups (i.e., the DS), defines the mechanical properties of the resulting HAGMbased hydrogels to a great extent.28 A DS of 90% was demonstrated by Bencherif and co-workers to provide the most stable hydrogels. A high degree of derivatization is considered to be crucial for 2PP processing. Previously described protocols only delivered materials with the maximum DS of 11−14%, which is not sufficient for 2PP fabrication.8,22 4.2. Mechanical Properties of the Photo-cross-linked Gels. In order to characterize the materials’ mechanical properties, rheological measurements were performed on a series of HAGM and HAGM−PEG hydrogels, respectively. The measurement of strain amplitude dependence of the storage and loss moduli (G′, G″) is well suitable to characterize linear viscoelastic behavior identifying the extent of the material’s linearity. Typically the rheological properties of a viscoelastic material are independent of strain up to a critical strain level γL. Following this method, linear-viscoelastic behavior was constituted for all tested materials as shown in Figure 3b, which is a main characteristic of HA to fulfill its mechanochemical function in the body (viscoelasticity of synovial fluid, vitreous humor in the eye).32 Obviously, the strain level γL depends mainly on the total dry hydrogel mass dissolved in the solvent, which is particularly significant for pure HAGM hydrogels. For the polymerized hydrogels as well as for the hybrid materials, average (G′) moduli were found (0.03− 0.45 MPa, data not shown) that correspond closely to that of naturally derived soft rubber (0.03−0.3 MPa).30 Thus, polymerization of HAGM leads to materials with a significantly increased mechanical load capacity in comparison to the mechanical properties of typical viscoelastic gels (50−5000 Pa). The combination of HAGM with different ratios of PEG allowed further modification of the mechanical properties showing the highest (G*) moduli among the series. Accordingly, materials with a high PEG content were demonstrated to have a narrow deformation range (Figure 3b) showing high stiffness (Figure 3c). This behavior corresponds to that of highly cross-linked polymers.31 Former studies provide strong arguments that elasticity and stiffness are 657

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on the 2PP fabrication process increasing the processing time to higher levels. Consequently, the copolymerization with synthetic prepolymers provides the opportunity to adapt 2PP processing parameters such as time, speed, power, etc. In summary, the utilization of CAD models opens access to predefined scaffold shapes and geometries, which can be created in a reproducible manner. Thus, this technique is ideal for the production of a series of identical hydrogel structures. In future, this hydrogel-based approach allows systematic studies of cell−scaffold interactions in a controlled three-dimensional environment.

adhesion, which also was demonstrated by several recent studies.22,28,35 Therefore, functionalization of the natural ECM-component HAGM with growth factors could perfectly adapt to the in vivo situation because the natural ECM also contains such biomolecules. In our work, human EGF, as model biomolecule, was covalently attached to HAGM using N-succinimidyl acrylate for linkage. Whether a physical or chemical modification with EGF was successful was analyzed by fibroblast growth under serum-free conditions. The exclusion of serum from the cell culture media is very important, since serum itself contains diverse growth factors. As a consequence, it could not clearly be determined whether the effects on cell growth correlate with the EGF-modified materials or serum. The problem is that the serum-free medium itself inhibits cell proliferation due to the missing growth factors and hormones and longer cultivation time can even induce apoptotic reactions. For fibroblasts, an average doubling time of about 35 h has been described.36 This means that after 96 h, a relative growth rate of about 400−500% would be expected under FCS conditions. On the control surface and on unmodified HAGM, a growth rate of less than 140% was reached under serum-free conditions (Figure 8). This supports the fact that no cell proliferation without growth factors can take place. When HAGM was modified with EGF, the growth potential was increased up to 177% in comparison to the unmodified value of only 68%. Therefore, it can be concluded that the functionalization was successful and EGF was bioactive. The disparity to cultivation with serum-containing medium can be explained by the fact that EGF, as one single component of serum, is not sufficient to stimulate cell growth in a normal way. Still other important components are missing to induce cell responses. Moreover, with the presented approach, further bone formation relevant biomolecules, such as bone morphogenetic proteins, could be attached to the hydrogel providing a specifically adapted microenvironment as shown by Shoichet and co-workers applying 2PP.37 However, the covalent attachment of methacrylated EGF to HA backbone might result in a decreased bioactivity compared with the natural EGF being released during the experiment. On the other side, for long-term application, a covalent attachment of growth factors might be favorable providing a constant depot for the time the tissue is being formed. In conclusion, it was proven that EGF functionalization can successfully be applied to fabricate innovative hydrogel scaffolds. 4.6. Scaffold Fabrication. Using computer-aided design models (CAD), we generated a diversity of possible structure geometries, demonstrating that tailor-made bioscaffolds can be produced to study cell-specific interactions on a series of identical samples. The fabrication of structures with both ringlike pores or grid constructs can be realized. Scaffolds with dimensions up to 2 mm were produced, which are well suited to handling, including cultivation with cells. Especially with regard to porosity requirements, 2PP is an ideal technique to produce structures with defined pore sizes and porosity (Table 1). Further, with this method, any other photosensitive material can be used to form hydrogel blend structures and hydrogel blends in situ applying 2PP or UV light, respectively. Thus, biocompatible poly(ethylene glycol) diacrylate (PEGDA) was used as a model component to prepare hybrid HAGM−PEG materials with different compositions affecting the biological and, primarily, the mechanical behavior of the hydrogels. The addition of this synthetic prepolymer shows a favorable effect

5. CONCLUSIONS HA is a major component of the human extracellular matrix and, therefore, an attractive starting material for bone tissue engineering. For the fabrication of mechanically stable scaffolds, a chemical functionalization of HA is required. For this purpose, we have prepared glycidyl methacrylated HA and combined HA−PEG hydrogels with different compositions in order to examine the biological and mechanical properties of the photo-cross-linked materials. Cell testing with osteoblasts confirmed the compatibility of the materials for the future usage as biomatrices for guided bone formation. The addition of different ratios of PEGDA was shown to modulate the mechanical properties of the generated gels without affecting the biocompatibility. To mediate cell responses, HA was covalently cross-linked with EGF and was compared with samples with encapsulated EGF. Applying two-photon polymerization, precisely defined three-dimensional HA and HA−PEG scaffolds with different geometries and pore sizes were successfully fabricated. Such structures provide promising prospects for cell investigations in a reproducible 3D organized hydrogel milieu.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Phone: +49 511 2788 233. Fax:+49 511 2788 100. Author Contributions

S.S.-W. and B.N.C. contributed equally. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The authors thank Dr. Jörg Fohrer (Institute for Organic Chemistry, Leibniz University Hannover) for providing the 1H NMR measurements. For the support given in terms of the rheology measurements, the authors thank Paul Haase-Aschoff from the Institute of Food Chemistry, Leibniz University Hannover. This work was supported by the DFG excellence cluster Rebirth “From Regenerative Biology to Reconstructive Therapy” and by the Low Saxony project “Biofabrication for NIFE”.

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REFERENCES

(1) Leclerc, A.; Tremblay, D.; Hadjiantoniou, S.; Bukoreshtliev, N. V.; Rogowski, J. L.; Godin, M.; Pelling, A. E. Biomaterials 2013, 34, 8097−8104. (2) Raimondi, M. T.; Eaton, S. M.; Laganà, M.; Aprile, V.; Nava, M. M.; Cerullo, G. M.; Osellame, R. Acta Biomater. 2013, 9, 4579−4584. (3) Drury, J. L.; Mooney, D. J. Biomaterials 2003, 24, 4337−4351. (4) Prehm, P. Biochem. J. 1983, 220, 597−600. (5) Fraser, T. C.; Laurent, J. R. E. FASEB J. 1992, 6, 2937−2944.

658

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Biomacromolecules

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(6) Alini, M.; Li, W.; Markovic, P.; Aebi, M.; Spiro, R. C.; Roughley, P. J. Spine 2003, 28, 446−453. (7) Kogan, G.; Soltes, L.; Stern, R.; Gemeiner, P. Biotechnol. Lett. 2007, 19, 17−25. (8) Leach, J. B.; Bivens, K. A.; Patrick, C. W., Jr.; Schmidt, C. E. Biotechnol. Bioeng. 2003, 82, 578−589. (9) Burdick, J.; Chung, C.; Jia, X.; Randolph, M. A.; Langer, R. Biomacromolecules 2005, 6, 386−391. (10) Xu, X.; Jha, A. K.; Harrington, D. A.; Farach-Carson, M. C.; Jia, X. Soft Matter 2012, 8, 3280−3294. (11) Koroleva, A.; Gill, A. A.; Ortega, I.; Haycock, J. W.; Schlie, S.; Gittard, S. D.; Chichkov, B.; Claeyssens, F. Biofabrication 2012, 4, No. 025005. (12) Koroleva, A.; Gittard, S. D.; Schlie, S.; Deiwick, A.; Jockenhoevel, S.; Chichkov, B. N. Biofabrication 2012, 4, No. 015001. (13) Li, J.; He, A.; Han, C. C.; Fang, D.; Hsiao, B. S.; Chu, B. Macromol. Rapid Commun. 2006, 27, 114−120. (14) Hunt, N. C.; Grover, L. M. Biotechnol. Lett. 2010, 32, 733−742. (15) Manjubala, I.; Scheler, S.; Bössert, J.; Jandt, K. D. Acta Biomater. 2006, 2, 75−84. (16) Farsari, M.; Chickov, B. N. Nat. Photonics 2009, 3, 450−452. (17) Malinauskas, M.; Farsari, M.; Piskarskas, A.; Juodkazis, S. Phys. Rep. 2013, 533, 1−31. (18) Togersen, J.; Qin, X.; Li, Z.; Ovsianikov, A.; Liska, R.; Stampfl, J. Adv. Funct. Mater. 2013, 23, 4542−4554. (19) Ovsianikov, A.; Malinauskas, M.; Schlie, S.; Chichkov, B.; Gittard, S.; Narayan, R.; Löbler, M.; Sternberg, K.; Schmitz, K.-P.; Haverich, A. Acta Biomater. 2011, 7, 967−974. (20) Ovsianikov, A.; Schlie, S.; Ngezahayo, A.; Haverich, A.; Chichkov, B. N. J. Tissue Eng. Regen. Med. 2007, 1, 443−449. (21) Ovsianikov, A.; Deiwick, A.; Van Vlierberghe, S.; Dubruel, P.; Möller, L.; Dräger, G.; Chichkov, B. Biomacromolecules 2011, 12, 851− 858. (22) Möller, A.; Krause, A.; Fahlmann, J.; Gruh, I.; Kirschning, A.; Dräger, G. Int. J. Artif. Organs 2011, 34, 93−102. (23) Kim, N. A.; Peppas, B. J. Drug Delivery Sci. Technol. 2006, 16, 11−18. (24) West, K.; Truong, J. L. Biomaterials 2002, 23, 4307−4314. (25) Fisher, J. P.; Dean, D.; Engel, P. S.; Mikos, A. G. Annu. Rev. Mater. Res. 2001, 31, 171−181. (26) El-Sayed, M. E. H.; Hoffman, A. S.; Stayton, P. S. Expert Opin. Biol. Ther. 2005, 5, 23−32. (27) Nicodemus, G. D.; Bryant, S. J. Tissue Eng., Part B 2008, 14, 149−165. (28) Bencherif, S. A.; Srinivasan, A.; Horkay, F.; Hollinger, J. O.; Matyjaszewski, K.; Washburn, N. R. Biomaterials 2008, 29, 1739− 1749. (29) Miyata, T.; Asami, N.; Uragami, T. Nature 1999, 399, 766−768. (30) Mezger, T. G. The Rheology Handbook: For Users of Rotational and Oscillatory Rheometers; Vincentz Network: Hannover, Germany, 2006; pp 80−81, 123. (31) Peng, T.; Yao, K.; Yuan, C.; Goosen, M. J. Polym. Sci. Part A: Polymer Chemistry 1994, 32, 591−596. (32) Necas, J.; Bartosikova, L.; Brauner, P.; Kolar, J. Vet. Med. 2008, 53, 397−411. (33) DeForest, C. A.; Anseth, K.S. Annu. Rev. Chem. Biomol. Eng. 2012, 3, 421−444. (34) Eng, D.; Caplan, M.; Preul, M.; Panitch, A. Acta Biomater. 2010, 6, 2407−2414. (35) Leach, J. B.; Bivens, K. A.; Collins, C. N.; Schmidt, C. E. J. Biomed. Mater. Res. 2004, 70A, 74−82. (36) Schlie-Wolter, S.; Ngezahayo, A.; Chichkov, B. Cell Res. 2013, 319, 1553−1561. (37) Wylie, R. G.; Ahsan, S.; Aizawa, Y.; Maxwell, K. L.; Morshead, C. M.; Shoichet, M. S. Nat. Mater. 2011, 10, 799−806.

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