Functional reconstruction of injured corpus cavernosa using 3D-printed hydrogel scaffolds seeded with HIF-1α-expressing stem cells

Injury of corpus cavernosa results in erectile dysfunction, but its treatment has been very difficult. Here we construct heparin-coated 3D-printed hydrogel scaffolds seeded with hypoxia inducible factor-1α (HIF-1α)-mutated muscle-derived stem cells (MDSCs) to develop bioengineered vascularized corpora. HIF-1α-mutated MDSCs significantly secrete various angiogenic factors in MDSCs regardless of hypoxia or normoxia. The biodegradable scaffolds, along with MDSCs, are implanted into corpus cavernosa defects in a rabbit model to show good histocompatibility with no immunological rejection, support vascularized tissue ingrowth, and promote neovascularisation to repair the defects. Evaluation of morphology, intracavernosal pressure, elasticity and shrinkage of repaired cavernous tissue prove that the bioengineered corpora scaffolds repair the defects and recover penile erectile and ejaculation function successfully. The function recovery restores the reproductive capability of the injured male rabbits. Our work demonstrates that the 3D-printed hydrogels with angiogenic cells hold great promise for penile reconstruction to restore reproductive capability of males.


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HAMA and GelMA was fabricated as follows. First of all, the natural cavernous sinus has a dense honeycomb-like porous structure. 4 Here, to simulate the overall shape and the internal honeycomb-like porous bionic structure of the corpora cavernosa tissue, we used computer-based 3D modeling software to design a 3D model (Supplementary Figure 21b). The designed model was stratified by 3D-Bioplotter TM software, and the printing model was imported to the control interface of the 3D-printing computer software. Then, loaded the 3D-printing ink into an extrusion unit using a single-channel needle (22G, 300 μm) after the nitrogen control system was turned on.
The temperature of the ink was controlled by a low-temperature control device (LTV-Dispense Head) on the 3D-Bioplotter TM (EnvisionTEC GmbH, Germany). The temperature of the printing platform was maintained at 4 °C. Other printing parameters, including fibre spacing, UV irradiation time, extrusion pressure and X-Y plotting speed, were set to 800 μm, 3 min, 8 × 10 4 Pa and 28 mm s -1 , respectively, to create straight and full printing fibres. The 3D-printed hydrogel porous scaffold was initially cured by hydrogen bonding between GelMA molecules at low temperature (4 °C) and then exposed to UV light (3 min, 10 mW cm -2 ) for further crosslinking. Preparation of acellular corporal collagen matrices (ACCM). ACCMs were obtained from the corpus cavernosum of the dead rabbits using our previously reported protocol. 5 Scanning electron microscopy (SEM). The 3D hydrogel scaffold was observed using a 3D rotational stereoscopic microscope (HiroX7700) to analyse the shape of the scaffold and the pore distribution after the swelling equilibrium was reached in PBS at 37 °C. The acellular corporal collagen matrices (ACCM), lyophilised normal corporal tissue, and lyophilised 3D hydrogel scaffold were observed by scanning electron microscopy (MERLIN, Carl Zeiss AG, Germany) after being fixed on a sample table and sprayed with gold for 90 s.
Compression testing and repeated compression stress. The compression testing and cyclic 4 compression testing of the 3D-printed hydrogel scaffold were performed on a universal material testing machine (Intron5967, Instron, USA) equipped with 500 N load sensors. All hydrogel scaffolds were soaked in PBS solution at 37 °C to reach equilibrium before experiments. The compression rate was 1 mm min -1 and at least three samples were tested in each group. The cyclic compression test of 3D-printed hydrogel scaffolds was carried out in the sweep of 0-40% strain for 20 cycles.

Layer-by-layer self-assembly (LBL) of heparin-coated 3D-printed hydrogel scaffolds.
Poly-L-lysine (PLL) and heparin were coated on the surface of the 3D printed hydrogel scaffold due to the electrostatic interactions between negatively charged heparin and positively charged PLL (Supplementary Figure 5a). The process is shown as follows. The 3D hydrogel scaffold was first immersed in a 1 mg mL -1 PLL solution for 30 min and then rinsed twice with ultra-pure water.
Next, the PLL-coated scaffold was immersed in a 10 mg mL -1 heparin solution for 30 min and then cleaned by ultra-pure water. The procedure was repeated 4 times to construct the heparin-coated 3D-printed hydrogel scaffold with the surface structure of PLL and heparin.
Zeta potential during the LBL process. For the convenience of testing the surface potential during the LBL process, the LBL process of the 3D hydrogel scaffold was described as follows.
The 3D-printing inks were coated on a rectangular glass sheet (2 cm x 1 cm) by spin coating (layer 0), and then layer 0 was exposed to UV light for crosslinking and curing before rinsed with deionised water 2 times. The zeta potential of layer 0 was measured using a solid-surface zeta potentiometer (Surpass) at pH = 7. Subsequently, after each modification of heparin and lysine (layer 1~9), the change values of each potential were recorded. X-ray photoelectron spectroscopy (XPS). The surface composition of the 3D hydrogel scaffolds and heparin-coated 3D-printed hydrogel scaffolds were analysed by a multifunction photoelectron spectrometer (ISIS-300, Oxford instruments). The vacuum degree was 2 x 10 -9 torr, and a mono-chromatic X-ray source of A1 Ka was used. The energy correction of the obtained XPS atlas was based on C1s (284.6 eV), and XPS PEAK41 software was used to fit the peaks of the characteristic elements (O, S and N) of heparin in the two groups of samples.
The release of heparin from the heparin-coated 3D-printed scaffolds. Due to the electrostatic interactions between toluidine blue O (TBO) and heparin, the surface density and 5 release period of heparin-coated 3D-printed scaffolds was characteried by TBO asasy. 6,7 The heparin-free scaffolds directly encapsulated with heparin served as a control group. A TBO solution was prepared by dissolving the 0.04 wt% TBO powder in 0.01 mol L -1 HCl/0.2 wt% NaCl aqueous solution. To establish a standard curve, 3 mL 0.04 wt% TBO solution was first added into 2 mL of a known concentration heparin solution (0.05 mg mL -1 , 0.1 mg mL -1 , 0.5 mg mL -1 and 1 mg mL -1 ) and incubated for 4 h at 37 o C on a 90 rpm shaker. Then, the heparin-TBO complex was spontaneously formed and precipitated in the mixture. The mixture was centrifuged at 8000 rpm for 10 min to remove the supernatant and the precipitate was carefully rinsed twice with HCl/NaCl solution (0.01 mol L -1 and 0.2 wt%, respectively). Finally, 5 mL ethanol/NaOH solution (80% and 0.2 wt%, respectively) was added to suspend the precipitate and the absorbance was recorded at 530 nm with a multifunctional microplate reader. There were five parallel samples in each group and the average value was used to make a standard curve (Supplementary Figure 6a).
To determine the heparin release curve, the heparin-coated 3D-printed scaffolds were immersed in 20 mL PBS at 37 o C on shaker (90 rpm) for 1, 4, 7, 10, 15, 25 and 30 days in a centrifugation tube.
The 2 mL release medium at each time point was collected and 2 mL fresh PBS was added in the release system. Subsequently, the collected medium was added into 3 mL TBO solution and incubated for 4 h in a shaker (37 o C, 90 rpm). Then, the quantitative characterization of heparin release was performed by using the same procedure as standard curve preparation. The released heparin density in the solution was calculated according to the standard curve, and the heparin release curve was plotted at last (Supplementary Figure 6b). Over five parallel samples were tested in each group.
Coagulation experiment. To further verify the anticoagulant effect of the scaffolds, the coagulation experiment was performed by soaking the heparin-coated scaffolds and the heparin-free scaffolds in fresh blood in the 37 ℃ shaker. The coagulation situations of different groups, including heparin-coated (heparin-coated scaffolds in the blood), heparin-free (heparin-free scaffolds in the blood) and blank (empty well with only blood) groups, were recorded at one-minute intervals.
In vitro degradation test. The degradation solution was prepared by dissolving 1 U mL -1 hyaluronidase and 0.001 U mL -1 protease in PBS. The cleaned heparin-free and heparin-coated 3D 6 hydrogel scaffolds were immersed in 10 mL degradation solution at 37 °C on a shaker (120 rpm) for 0, 6, 20, 35, 50, 65, 80 and 100 days in a centrifuge tube. Recorded the weight of the scaffolds after lyophilisation at each time point.

Subcutaneous implantation in nude mice in vivo.
To evaluate the synergistic function of the cells and material to promote vascularisation during the tissue repair process, normal MDSCs, vector MDSCs and mHIF-1α MDSCs at 10 6 cells mL -1 were seeded on heparin-free and heparin-coated 3D scaffold materials (5 mm × 5 mm × 2 mm). Then, the cell-loaded scaffolds were cultured in vitro at 37℃ for 7 days to fully adhere and spread on the surface of the material, and cell growth was observed using laser confocal microscopy. Seven-week-old nude mice were used to establish a subcutaneous implantation model in this study and divided into 5 groups (cell-free and heparin-free scaffolds, MDSCs-loaded heparin-free scaffolds, vector cell-loaded heparin-free scaffolds, mHIF-1α cell-loaded heparin-free scaffolds, and mHIF-1α cell-loaded heparin-coated scaffolds). All animal procedures were approved by the Animal Ethical and Two-photon imaging of the blood vessels. Vascularity could be visualised in the materials through two-photon imaging technology. Primarily, the mice were anaesthetised with 1% pentobarbital (50 mg kg -1 ) by intraperitoneal injection and fixed on the stage of a two-photon microscope (Leica, DM6000, German). To visualise the vasculature, 0.2 mL 1% fluorescein isothiocyanate (FITC, Sigma, Germany) saline solution was injected intravenously into the tail of mice. Then, the skin and fascia were carefully separated over the surface of the material, and a 2×2 mm 2 window was prepared before imaging. Two-photon images of blood vessels was obtained at a wavelength of 800 nm with a laser scanning system (Coherent, Santa Clara, CA, USA). Whole images of the X-Y-Z stacks (1024 × 1024 pixels, 2 mm resolution) were reached to 300 um below the scaffold surface to show the distribution and density of vasculature.
Bionic mechanics test of the scaffolds. After 30 days of implantation, the heparin-coated 3D hydrogel scaffold implants in the nude mice (7-week-old) were obtained (Supplementary Figure   16a). Then, a compression test was employed to assess the mechanical properties of the samples before and after implantation. The compression rate was 1 mm min -1 .   Table 1. Proportion of surface elements on heparin-free and heparin-coated 3D hydrogel scaffolds calculated via XPS.