2.1.

Photosensitive Polymer Materials

Three different hybrid organic-inorganic polymer materials were used in our experiments. Two of them are commercially available UV curable materials Ormoclear and Ormocore b59 produced for photo-lithographic applications in microoptics (Micro Resist Technology GmbH). These materials belong to the ORMOCER class hybrid polymers and they are produced using sol-gel process from liquid precursors. They include inorganic oxide units and cross-linkable organic moieties such as urethane-and thioether (meth)-acrylate alkoxysilanes, and contain strong covalent bonds between the inorganic and organic components.⁴¹ These materials were used as purchased without any further modifications (photoinitiatior is already added by the suppliers). These materials have been tested for scaffold production with DLW and biocompatibility before.⁴² Another material is a two composite sol-gel based on zirconium and silicon oxides (SZ2080). The preparation of it is described elsewhere.⁴³ It contains inorganic silicon alkoxide and zirconium alkoxide groups which enhance materials mechanical stability, while organic part of SZ2080 contains polymerizable methacrylate moieties. This low shrinkage photopolymer was initially used to fabricate photonic elements,⁴⁴ but also found its applications in TE.²⁴

SZ2080 was photosensitized by adding 1 to 2 wt. % of 2-Benzyl-2-dimethylamino-1-(4-morpholinophenyl)buta-none-1 photoinitiator (also known as Irgacure 369) (Sigma-Aldrich GmbH). This commercial photoinitiator was chosen because it has already been proved for suitability for high quality and throughput fabrication via DLW.^45,46 Although common commercial photoinitiators may suffer from low two-photon absorption efficiency⁴⁷ compared to the novel custom-made chromophores,^48,49 yet preparation of organic molecules with large two-photon absorption cross sections is rather complicated and was not involved in this research.

Before fabrication samples were prepared by drop-casting the materials on the microscope cover glass substrates. After laser processing the photosensitive resins exposed to light underwent polymerization and became insoluble in the developer [Fig. 1(a)]. Samples were treated with the appropriate organic solvent in order to wash out unexposed material. Polymerized structures sustained during the development process. In this way, the free-standing structures were fabricated on a glass substrate. Scanning Electron Microscopy (SEM) and optical profilometry were applied to evaluate the micro-structured scaffolds.

Fig. 1

(a) Fabrication steps: (I) photopolymerization reaction is initiated in the focal spot of the beam, (II) polymeric structure is directly written by moving the sample in regard to the focal spot, (III) organic developer washes unexposed material (IV) 3D free-standing scaffold is obtained on a glass substrate; (b) DLW fabrication setup: ATN-attenuator, NC-second harmonic non-linear crystal, M1-M2-mirrors, L1-L2-telescope, DM-dichroic mirror, OL-objective lens, BS-beam splitter, CMOS-camera.

2.2.

Direct Laser Writing Setup and Scaffold Fabrication

The schematics of the DLW system used in this work is depicted in Fig. 1(b). High peak power femtosecond Yb:KGW laser amplifier (Pharos, Light Conversion Co. Ltd.) was used as irradiation source. Its parameters are: 300 fs pulse duration, 200 kHz repetition rate, 1030 nm central wavelength (515 nm second harmonic was used in experiment). 1 to 10 mW of average laser power was used which corresponds to $0.32-3.2\mathrm{TW}/{\mathrm{cm}}^2$ (values are calculated assuming that $10\times 0.3\mathrm{NA}$ objective was used, which was the most practical for the fabrication of the scaffolds in our experiment).

Irradiation was calculated using $I_p=\frac{T{E}_p}{{\tau }_p\pi \omega }$,⁵⁰ where T is the objective’s transmittance for 515 nm, $E_p=P/f$ is the pulse energy, ${\tau }_p$ is the pulse duration and ${\omega }_p=\frac{0.61\lambda }{\mathrm{NA}}$ is the beam waist radius. For the highest translation velocity achieved we calculated the cumulative dose of exposure $D_{ac}=D_p\times N_p$, where $D_p$ is the dose per pulse ($D_p=I_p{\tau }_p$) and $N_p$ is the number of pulses per focal spot. The estimated cumulative dose was $4.6\mathrm{J}/{\mathrm{cm}}^2$ showing an efficient energy use for polymerization reaction when using fs pulses. All necessary values used for calculations and calculated quantities are given in Table 1.

Table 1

Fabrication parameters (NA-objective’s numerical aperture, λ-laser radiation wavelength, τp-pulse duration, f-pulse repetition rate, T-objective’s transmittance for 515 nm, υ-sample scanning velocity, P-average laser power, ω-beam waist radius, Ep-pulse energy, Ip-irradiation, Np-number of pulses per focal spot, Dp-dose per pulse, Dac-cumulative dose).

The laser beam was guided through an optical system [Fig. 1(b)] to a high numerical aperture objective and focused to a volume of photopolymer. The sample was mounted on high speed and wide working area positioning system which consisted of linear motor driven stages (Aerotech, Inc.): XY-ALS130 to 100, Z-ALS130 to 50. These stages ensure an overall traveling range of 100 mm in $X$ and $Y$ directions and 50 mm in $Z$ direction and support the scanning velocity of up to $300\mathrm{mm}/\mathrm{s}$ while keeping positioning resolution up to 10 nm. Upon irradiation the monomers underwent transition from liquid to solid (or from gel to solid) which resulted in the change of the refractive index. It enabled wide-field transmission microscopy to be used for monitoring the manufacturing process in real time. A microscope was built by adding its main components to the system: a source of red light provided by a LED, a CMOS camera (mvBlueFOX-M102G, Matrix Vision GmbH) and a video screen. The ability to image photostructuring while performing DLW is an important feature for a successful fabrication process. Control of all equipment was automated via custom-made software 3DPoli specially designed for DLW applications.* By moving the sample three-dimensionally, the position of the laser focus was being changed inside the resin and this enabled writing of complex 3D structures. Structures can be imported from CAD files or programmed directly. This DLW system was tested for structuring in various photosensitive materials at large scale. The ability to scale up and speed up the fabrication was ensured by changing the laser beam focusing objectives in the range from $100\times \mathrm{NA}=1.4\text{\hspace{0.17em}to}10\times \mathrm{NA}=0.3$, thus at the sacrifice of the resolution from 200 nm to 4 μ m.⁵¹

2.3.

Biocompatibility Tests In Vitro and In Vivo

Biocompatibility of the micro-structured materials was assessed in vitro using stem cell culture and in vivo using laboratory rats (Lithuanian State Food and Veterinary Office license for use of laboratory animals for scientific research work No. 0212, 2011 to 02-04). Laboratory rats were taken care of according to the Lithuanian Law on animal care, housing and use (No. VIII-500, 1997 to 11-06).

For the studies in  vitro, a primary myogenic stem cell line derived from adult rabbit skeletal muscle was established.⁵² Before cell seeding, the specimens were sterilized with 70% ethanol for 24 h, and dried on a clean bench under UV irradiation. The cells were grown in Iscoves modified Dulbeccos medium (Sigma-Aldrich GmbH) supplemented with 10% fetal calf serum (Biological Industries Ltd.) and antibiotics (penicillin, 100 U/ml, streptomicin, $100\mu \text{g}/\text{ml}$; Biological Industries Ltd.), at 37°C and 5% ${\mathrm{CO}}_2$. For the evaluation of biocompatibility, the cells ($6\times {10}^4/\mathrm{ml}$) were grown for 48 h on the tested micro-structured polymeric scaffolds which were placed inside the wells of a 24-well plate (Orange Scientific). Cell viability was examined using differentiating staining with acridine orange (AO, $100\mu \text{g}/\mathrm{ml}$, Molecular Probes Inc.) and etidium bromide (EB, $100\mu \text{g}/\mathrm{ml}$, Sigma-Aldrich GmbH).⁵³ 4 μl of AO/EB dye mixture (1∶1) was added to 100 μl of growth medium on the cell monolayer. Fluorescent microscopy (Nikon Instruments Inc.) analysis allows to distinguish living (green) and dead (orange) cells in cell population stained this way. On the basis of cell membrane integrity, AO binds to DNA of viable cells and stains them green, as well, EB bounds to the DNA of nonviable cells and stain them in orange.

In vivo study was performed by using healthy male Wistar rats, weighting 250 to 300 g (Vivarium of the Institute of Biochemistry, Vilnius University). Prior to the surgery, the polymeric structures out of Ormocore b59 and SZ2080 were washed in ethanol ($3\times 24\mathrm{h}$) and dried under UV ($2\times 15\min$, both sides). The rats were anesthetized by intramuscular injection of ketamine hydrochloride ($100\mathrm{mg}/\mathrm{ml}$; $40\mathrm{mg}/\mathrm{kg}$) (Vetoquinol Biowet, Poland) and xylazine hydrochloride (2%; $5\mathrm{mg}/\mathrm{kg}$) (Bela-pharm, Germany). The lumbar area of the rat was prepared aseptically for surgery, longitudinal incisions of approximately 1 cm were made in the skin near the paravertebral back muscle where two samples of each polymer were inserted. The incisions were closed with interrupted sutures of Polysorb 4 to 0 suture (Syneture). After 3 weeks, the animals were euthanized, each implant was removed and a histological analysis of the surrounding tissues was performed in the National Center of Pathology (Vilnius, Lithuania). Biocompatibility of the tested polymeric samples was evaluated judging by the microscopic views of the histologic specimens of surrounding tissues stained with hematoxylin and eosin (H&E).

3.1.

Fabrication of 3d Scaffolds

One of the most important advantages of DLW compared to alternative above mentioned technologies is the precise control of structuring resolution, which allows fabricating precise objects with almost no geometrical restraints. Spatial resolution can be flexibly tuned by varying laser output power and translation velocity of the sample, by replacing focusing objectives or by altering sensitivity of the material itself (changing the concentration of the used photoinitiator).

To the best of our knowledge up to date, the structure shown in Fig. 2 is the largest 3D micro-structured scaffold produced by femtosecond direct laser writing in polymers.⁵⁴

Fig. 2

(a) A photo of $9\times 9\times 2{\mathrm{mm}}^3$ disc shape scaffold out of SZ2080 polymer; (b) Inset shows SEM image of internal structure of the disc shape scaffold. The pore size is $50\times 50\times 100{\mu \text{m}}^3$, and the general porosity is 60%. $5\mathrm{mm}/\mathrm{s}$ sample translation velocity and $10\times 0.3\mathrm{NA}$ objective were used; (c) A photo of an artificial blood vessel scaffold out of Ormoclear polymer with outer diameter of 3 mm, internal diameter 1.5 mm, length of 9 mm and a 45° cut edge. The pore size is $10\times 10\times 100{\mu \text{m}}^3$, and the general porosity is 67%. $10\mathrm{mm}/\mathrm{s}$ sample translation velocity and $10\times 0.3\mathrm{NA}$ objective lens were used; (d) SEM image of an artificial blood vessel scaffold fabricated out of Ormoclear polymer with a 1.5 mm outer diameter and a 0.75 mm internal diameter with a length of 2 mm. The pore size is $50\times 50\times 100{\mu \text{m}}^3$, and the general porosity is 50%. $2\mathrm{mm}/\mathrm{s}$ sample translation velocity and $10\times 0.3\mathrm{NA}$ objective lens were used.

3.2.

Biocompatibility Test of Hybrid Materials

The biocompatibility of micro-structured polymeric scaffolds fabricated out of Ormocore b59 and SZ2080 was evaluated in vitro using rabbit muscle-derived myogenic cells which were grown on the tested materials. Polystyrene was used as a control surface. The results demonstrate that the viability of cells grown on polymeric scaffolds is comparable to that on control surfaces. In both cases, less than ten percent of dead cells were found (Fig. 3).

Fig. 3

Living rabbit muscle-derived myogenic stem cells growing in vitro on the micro-structured Ormocore b59 (a) and SZ2080 (b) polymeric surfaces. (c) The distribution of viable and non-viable cells growing on tested surfaces according to AO/EB staining. Cells were categorized as follows: V-viable with non-fragmented nuclei (bright green chromatin); VA-viable with fragmented nuclei (bright green chromatin with organized structure); N necrotic (red non-fragmented nuclei); NVA-non-viable apoptotic with fragmented nuclei (bright orange chromatin that is highly condensed or fragmented). Chromatin-free cells lost their DNA content entirely and exhibited weak green-orange staining.⁵³

For the biocompatibility studies in vivo, after three weeks of implantation, histological analysis of tissues surrounding the implant was performed. It is known that the presence of the implant changes the healing process. A reaction of biological tissue to any foreign material is called the Foreign-Body Reaction (FBR). Foreign bodies can be inert or irritating, some of them can be toxic and cause inflammation as well as scarring. The normal FBR consists of macrophages and foreign-body giant cells at the surface of the implant with subjacent fibroblastic proliferation and collagen deposition, and capillary formation. Macrophages play a pivotal role in the response of tissue to implants.⁵⁵ As revealed in the Fig. 4, a mild FBR, characterized by the sporadic presence of macrophages and lymphocytes, is observed in the tested slides. Taking into account that such tissue reaction is registered three weeks after the implantation, the results are considered to be positive and demonstrate that tested materials were non-cytotoxic and biocompatible, showing them to be suitable for biomedical practice.

Fig. 4

Biocompatibility tests in vivo: (a) tissue reaction to the nonstructured Ormoclear and (b) to micro-structured SZ2080; mild FBR, characterized by a small number of macrophages and lymphocytes around the implant location is observed in the tested slides.

Comparing test results in vitro and in vivo we conclude that studied polymeric structures are biocompatible and suitable for tissue engineering applications.