A strategy for rapid and facile fabrication of controlled, layered blood vessel-like structures

Abstract

We develop a rapid and facile method to fabricate tubular scaffolds by a single-step rolling operation. With the aid of fibrin medical glue and a smooth expanded polytetrafluoroethylene (ePTFE) mandrel, we can wrap a piece of flat thin film into a three-dimensional (3D), multi-layered tubular structure with well-controlled diameter, wall thickness, and mechanical strength within 10 min. By patterning different cells in a pre-designed area on the film, after rolling, we can obtain blood vessel-mimicking tissues with well-arranged, multilayered 3D architectures within 70 min. Our strategy provides an excellent platform to rapidly fabricate tubular scaffolds essentially with no equipment and straightforward manipulations.

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1. Introduction

Blood vessels generally constitute three layers: intima (inner layer), media (middle layer), and adventitia (outer layer), each of which contains a specific cell type: endothelial cells (ECs), smooth muscle cells (SMCs), and fibroblasts (FBs), respectively.¹ These cells are distributed in a well-organised architecture so that the blood vessel performs its special functions, such as regulation of the blood biochemical environment, control of blood flow distribution, and maintenance of vessel elasticity.¹ Reconstruction of blood vessel-mimicking structures in vitro is of great significance for regenerative medicine and drug screening.^2–8 Development of a rapid and facile process to replicate these structures is a critical step for their further assessment and application. At present, although techniques for this purpose have been widely developed, there is still large room to develop a more rapid and facile method to fabricate blood vessel-like structures.

The efforts on blood vessel-like structure reconstruction can be dated back to 1980, where different layers of cell-matrix containing SMCs and FBs were cast into annular molds before the injection of ECs to form an endothelium.⁹ This structure possesses many characteristics of a real blood vessel, such as layered structures and electrophysiological responses. However, it needs at least one month to form the whole structure. After that, various strategies have been employed to create blood vessel-like structures. In a method developed by our group, stress-induced rolling membrane (SIRM) technique and cell patterning were combined together to create blood vessel-like tubular structures using stretched PDMS as the substrate material.^2,3,10 This method inspires us that cell patterning can be applied on flat films to facilitate the subsequent cell arrangement. Although the method greatly reduces the fabrication period, the limitations of this method are obvious. Firstly, PDMS is a non-biodegradable material, which does not benefit future tissue engineering use; meanwhile, the materials option has confined room due to the requirement of elastic layer that is a requisite for the fabrication process. Secondly, control of the vessel diameter is not quite precise because the elasticity of stretched PDMS films will impact the control process. Thirdly, the different layers are not bonded with each other. This may cause liquid leakage from the gaps under blood pressure. We further found that these challenges could be well addressed by introducing the technique of rolling biocompatible materials around a mandrel.

Several groups have reported the fabrication of vascular graft rolling around a mandrel. However, these grafts normally need prolonged periods to form vessel-like structures, and/or cannot form well integrated wall in a short time, and/or need complicated steps and manipulations to form the blood vessel-like structure with three kinds of vascular cells. For example, stacking different layers of SMCs and FBs with thick cell matrix around a PTFE mandrel could construct a fully biological blood vessel-like structure.¹¹ In this method, the SMC- and FB-containing cell sheets stimulated by sodium ascorbate needed around one month to be harvested. In the fabricating process, the first layer was a dehydrated FB-containing sheet; the second SMC-containing layer was wrapped around the first layer and cultured in a bioreactor for a week to increase the layer fusion; the third FB-containing layer was added later and cultured for at least 7 weeks to fully fuse the different layers; after removal of the mandrel, ECs were seeded on the lumen of the tubular tissue and grew for around one week. This method takes at least 3 months to complete the fabrication. The manipulation and layer integrity formation are also very complex. An improved fabrication method allows a single step rolling method.¹² After the preparation of cell sheets by culturing SMC and FB adjacently in the same dish with sodium ascorbate stimulation, the cell sheets were wrapped around a PTFE mandrel and kept for 2 weeks before removal of the mandrel. The time for the whole process can be reduced to around 1.5 months. Although quicker than the previous method, the fabrication and fusion time is still rather long and the manipulation is not simple enough. In other strategies, the cell sheets can also be produced on a thermoresponsive culture surface and then used for vessel fabrication.¹³ The SMC sheets were detached from the poly(N-isopropylacrylamide) (PNIPAm) surface after a 4 day culture, put on a PCL ES film to form a polymer–cell hybrid film. Then the film was wrapped around a PTFE mandrel and kept in culture medium for several days to form a constructed blood vessel. We can find that this approach cannot form three-layered vessel structures and the layer wall is not tightly integrated. Electrospun PCL film seeded with SMC cells can be wrapped around a stainless steel mandrel to form blood vessel media.¹⁴ However, the problem remains that this approach cannot form three-layered vessel structures and the layer wall is separated but only fastened by suture lassos at its ends. Hydrogel seeded with three kinds of vascular cells can be transferred by water printing on an electrospun PLGA film and the film can be wrapped around a mandrel to form multilayered cell-laden vessels.¹⁵ This strategy can rapidly form three-layered blood vessel-like structures. But the different layers are also not stably fixed. Moreover, the control of cell distribution in this approach is not very precise because the cells on the film would be left proliferation for several days before rolling.

After collective consideration of these approaches, we developed a rapid and facile strategy to fabricate blood vessel-like structures by a single step rolling operation. In our strategy, biodegradable electrospun films were employed as the major substrate material for their good biocompatibility; the biomedical glue (fibrin glue in specific in this study) was used as the bonding material to significantly accelerate the structure formation and increase the integrity of different layers, which is enlightened from the structure of a foam tape (Fig. 1A); polydimethylsiloxane (PDMS) chambers were employed to confine different cells in predesigned areas on the film to precisely control the cell distribution in formed tubes; cell suspensions rather than cell sheets were used to eliminate the time-consuming cell sheet formation and shorten the cell seeding and adhesion period. By this means, we could obtain scaffolds with controllably distributed cells in 3D architecture within 70 minutes. We could also conveniently achieve flexible choice of materials, precisely defined diameter, and tunable layer numbers when fabricating these structures. In this paper, we will concentrate on introducing the fabrication process and discussing its features and potential use.

Fig. 1 (A) The structure of the foam tape. The layers are rolled around a mandrel and maintained the integrity by glue. (B) The brief gelation mechanism of the fibrin glue.

2. Experimental section

2.1. Materials

Foam tapes, ePTFE tubes, and iron wire meshes were purchased from Huawei Wujing Shop in Beijing. The ePTFE tubes are rigid and hard to deform in our fabrication with gentle force. For a 3 mm outer diameter tube, it has a wall thickness of 1 mm. Polylactic-co-glycolic acid 50 : 50 (PLGA 50 : 50) (m_PLA : m_PGA = 50 : 50) and PLGA 75 : 25 (m_PLA : m_PGA = 75 : 25) polymers were of pharmaceutical grade and purchased from Lakeshorebiomaterials Co., Ltd. Poly(ε-caprolactone) (PCL) and fibronectin (FN) were obtained from Sigma. The fibrin medical adhesive (Porcine Fibrin Sealant Kit) was bought from Puji Medical Technology Development Co., Ltd (Hangzhou, China). The adhesive can be directly used in hospital surgery. Component A contains 4 mg ml⁻¹ fibrinogen and 9 mg ml⁻¹ sodium chloride, and component B contains 450 IU ml⁻¹ thrombin and 4.44 mg ml⁻¹ calcium chloride. In the fabrication process with a maximal period of ten minutes, amounts of 10 μl cm⁻² were used for each sample. Other reagents were all of analytical grade obtained from Beijing Chemical Factory. The ePTFE tubes were sterilised by soaking in 75% ethanol for 1 h before use.

2.2. Bacterial cellulose film preparation

For bio-synthesis of bacterial cellulose (BC), the bacterium, gluconacetobacter xylinum (ATCC23767), was cultured in a Hestrin and Schramm (HS) medium.¹⁶ The medium is composed of 2% (wt) glucose, 0.5% (wt) yeast extract, 0.5% (wt) peptone, 0.27% (wt) disodium phosphate (Na₂HPO₄), and 0.15% (wt) citric acid. After incubating in sterilised medium statically for 6 days at 26 °C, the BC membranes were dipped into distilled water for 2–3 days, and steamed by boiling in 1% (wt) NaOH solution for 30 min. Then, the films were washed with distilled water and high-purity water several times until the pH reached 7.0. The films were sterilised for 20 min at 121 °C, then air dried for 1–2 days, and left for use.

2.3. PLGA and PCL ES film preparation

Acetone/N,N-dimethylformamide (DMF) blend solution was mixed with a ratio of 2 : 1 (w/w). PLGA (m_PLA : m_PGA = 75 : 25 or 50 : 50, M_w ≈ 114 kDa) was dissolved in the blend solution at 20 wt%. The electrospinning (ES) was performed at 11.5 kV voltage with the spinning distance of 10 cm. PCL partials were dissolved in the blend of CH₂Cl₂ and N,N-dimethylformamide (DMF) with a ratio of 3 : 1 (w/w) to obtain a concentration of 20 wt%. The ES was performed at 12 kV voltage with the spinning distance of 10 cm. The films are sterilised by cobalt ray radiation (10 kGy) before use. All films to be seeded with cells were incubated with FN (20 μg ml⁻¹ in PBS) at 37 °C for 1 h to enhance cell adhesion and washed with PBS once before use.

2.4. PDMS and PMMA substrate fabrication

Patterned PDMS substrates were made via soft lithography. Briefly, SU-8 photoresist was spin-coated onto clean wafers. The wafers were then exposed to UV light through the transparency mask printed. The printed mask patterns were designed according to the substrate we needed. PDMS pre-polymer and catalyst (Silpot 184 and Catalyst Silpot 184, Dow Corning, US) were mixed at a ratio of 10 : 1 and cured against the patterned wafers at 80 °C for 2 h. PDMS substrates were then carefully peeled off the wafer surface. Polymethyl methacrylate (PMMA) substrates were fabricated by digitally controlled micromachining.

2.5. PLGA cast film preparation

PLGA was dissolved in dichloromethane (CH₂Cl₂) at a concentration of 5% (w/w). PLGA solution was cast on a polydimethylsiloxane (PDMS) substrate with microstructures. After evaporation in a fume hood for 2 h, the film was put into a vacuum drying oven for 12 h. Then, the film was peeled off from the substrate and sterilised by cobalt ray radiation (10 kGy).

2.6. PDMS chamber fabrication

PDMS pre-polymer and catalyst were mixed at a ratio of 10 : 1 and cured against the patterned polymethyl methacrylate (PMMA) substrate at 80 °C for 2 h. PDMS chambers were then carefully peeled off the substrate surface. The extra PDMS on top of the chamber was cut off using a knife to facilitate liquid injection.

2.7. Cell culture, staining, and seeding

Three kinds of cell lines: HUVECs (ATCC, US), HASMCs (ATCC, US), and HAFs (Science Cell, US) were cultured in DMEM (Invitrogen, US) containing 10% fetal bovine serum (FBS, Invitrogen, US), 1% penicillin–streptomycin (PS, Invitrogen, US), 1% Gluta-max (Invitrogen, US) at 37 °C, 5% CO₂. Every passage of cells was conducted with 0.25% trypsin with 0.02% ethylenediamine tetraacetic acid (EDTA, Invitrogen, US) every 2–3 days. For the scaffolds used to visualize the cell distribution, before seeding into PDMS chambers, HUVECs, SMCs, and HAFs were stained with dyes with different colours by Cell Tracker Orange (Invitrogen, US), Cell Tracker Green (Invitrogen, US), and Celltracker Deep Red (Invitrogen, US), respectively. We then collected cells and delivered them into the PDMS chambers at a density of 2 × 10⁴ cm⁻². Before pipetting the cell suspension into designated channels, PDMS chambers had been placed on the PLGA film whose spacer ridges between channels had been sealed by fibrin glue to avoid liquid leakage. For the scaffolds used to evaluate cell viability after culture, the cell staining step is not necessary.

2.8. Fabrication of layered scaffolds

To fabricate the layered scaffolds without cells, the films (including PLGA ES films, PCL ES films, PLGA cast films, and BC films) were cut into rectangles with proper sizes, coated with component A of the fibrin glue on one side and component B on the other, and rolled around an ePTFE mandrel with proper outer diameters by hand at a speed of 2–4 cm min⁻¹. When rolling, the two components of the fibrin glue would react with each other and bond the layers. The mandrel was gently extracted and the scaffold was ready. To fabricate the layered scaffolds with cells, cells were patterned according to the procedure above described. The PDMS chamber was peeled off and the scaffolds were rolled up just the same as those without cells.

2.9. Measurement of film thickness

The thickness of films was measured by an electronic digital micrometer (Guanglu, China). 5 positions on each film were selected randomly, and the average of the 5 values was determined as the thickness of the film. For a patch of 10 films to be used, the thickness was measured, and expressed as mean ± standard deviation (sd).

2.10. Measurement of mechanical properties

Suture retention and burst pressure were performed according to the ANSI guidelines (ANSI/AAMI/ISO 7198:1998/2001/(R)2010 Cardiovascular implants: Tubular vascular prostheses). Suture retention: a 1 cm long graft was clamped at one end onto a dynamic mechanical analysis machine (DMA, Q800, TA instrument, US) and a single bite suture (7-0 prolene suture, Ethicon, UK) was placed 2 mm from the edge of the other end. A constant pulling rate of 1 mm min⁻¹ was applied until the suture was pulled out. The maximum force of pulling was recorded as the suture retention. Burst pressure: one end of the graft (2 cm) was hermetically clamped, and the other one end of the graft was hermetically connected to a syringe on a syringe pump (PHD ULTRA, Harvard, US). A constant filling rate of 0.1 ml min⁻¹ was applied. The pressure when the tube burst was tested by a pressure gage as the burst pressure (AZ 82100 and AZ 8205, Taiwan, China). All tests were repeated for 3 times.

2.11. Cell distribution and microstructure observation

To observe cell distribution in the cell-seeded scaffold, the scaffold was embedded in PDMS and sectioned into thin slice as previously described.² Briefly, the tube was immersed into liquid PDMS, and bubbles were removed from one end of the tubes with a pipette. After curing the liquid PDMS (2–3 h, 50 °C), the PDMS was cut into thin slices. The layered structure was observed under confocal microscopy (Zeiss LSM 710, Germany). The emission/maximum excitation wavelengths of three kinds of stains are Cell Tracker Green: 488/517 nm, Cell Tracker Orange: 543/565 nm, and Celltracker Deep Red: 633/650 nm.

The microstructures of tubes, films, and fibrin glue fibers were observed by scanning electron microscopy (SEM, Hitachi S-3400N, Japan) and/or an upright light microscope (Leica DMI 6000B, Germany).

2.12. Cell viability test

To test the impact of dissolved glue components on cell viability, we dissolve 20% (v/v) component A, 20% (v/v) component B, and 20% (v/v) both component A and B in PBS, and incubated with HUVECs seeded in 96-well culture plate (1 × 10⁴ cells per well) under ambient temperature for 10 min, 30 min, and 120 min. The cells were then washed with PBS and stained with the LIVE/DEAD kit (Invitrogen, US). The viability was also tested at the concentration of the test components up to 60%. The image was observed with confocal microscopy. The cells incubated with PBS and DMEM were used as the control.

To test the impact of glue components on cell seeded on PLGA film, we cultured HUVECs on PLGA films at a density of 5 × 10⁴ cm⁻², and coated the cells with component A, component B, and a mixture of component A and B. The films were incubated under ambient temperature for 10 min, 30 min, and 120 min. The cells were then washed with PBS and stained with the LIVE/DEAD kit. The image was observed via confocal microscopy. The films incubated with PBS and DMEM were used as the control.

To test the cell viability in the scaffold after culture, we collected the HUVECs, HASMCs, and HAFs without staining, and delivered them into the PDMS chambers at a density of 2 × 10⁴ cm⁻². After static culture in DMEM medium at 37 °C, 5% CO₂ for 3 days, we unrolled the scaffold, washed the cells on the film with PBS, and stained with the LIVE/DEAD kit. The image was observed via confocal microscopy.

3. Results and discussion

3.1. The inspiration and the selection of materials

A significant inspiration of our method to involve biomedical glue is from the structure of the foam tape, a widely-used material in everyday life and industry. The tape is made up of rolled foam films around a mandrel and the layers of films are stuck by glue (Fig. 1A). The glue plays a decisive role in maintaining the integrity of the foam tape. Considering biocompatibility and processing facility, we choose fibrin medical adhesive, a kind of biological glue with two components, as the glue in our scaffolds. Fibrin medical adhesive has been widely used in surgery.¹⁷ It is low-immunogenic and highly bioabsorbable, whose stickiness is strong enough to allow the bonded tissues to resist physiological blood pressure. Component A contains fibrinogen and factor XIII dissolved in NaCl aqueous solution, and component B contains thrombin dissolved in CaCl₂ aqueous solution. When the two components contact each other, they form the fibrin gel within seconds (Fig. S1†) and the gelling mimics the last step of blood coagulation (Fig. 1B and S2†). We choose polylactic-co-glycolic acid (PLGA), an FDA-approved biodegradable material, as the material for the substrate film.¹⁸ The film is prepared with two methods: one is electrospinning (ES), and the other is casting. For ES method, the film thickness can be controlled by modulating the PLGA solution concentration, spinning voltage, and ES time (Fig. S3†). For casting, the film thickness can be controlled by varying solution concentration and volume. Other films of different materials (such as bacterial cellulose (BC)) or with microstructures are also used to demonstrate the flexibility and universality of our method. PDMS chambers with predesigned geometries are used to pattern cells. For the mandrel, we use expanded polytetrafluoroethylene (ePTFE) tubes with desired outer diameters (Fig. S4†). These tubes are smooth, rigid, with good cell compatibility, as well as easy to obtain and produce.¹⁹

3.2. Fabrication of cell-free scaffolds

The fabrication of the layered tube without cells starts from preparing a rectangular thin film with proper size. The two sides of the thin membrane are coated with the two components of fibrin adhesive separately (Fig. 2A). Subsequently, the film is wrapped around a smooth ePTFE mandrel manually. The consistent but gentle force by fingers (wearing sterile gloves) is applied to roll the mandrel and keep the film in touch with each other. The rolling speed is controlled around 2–4 cm min⁻¹ to allow sufficient reaction but not drying of the glue. When wrapping, the two components of the fibrin glue contact, form gel within 3 minutes, and bond the film layers. After the mandrel is gently extracted, the tube with fixed layered walls can be stably maintained by the glue, which becomes a self-standing hollow tubular structure. The microstructure of the reacted glue is shown in Fig. S5.† The unreacted residual glue components coated on the film surface will then be dissolved in PBS or culture medium and result in no harm (Fig. S6†). Fig. 2B and C show a PLGA 50 : 50 ES tube with 3 mm inner diameter and the SEM imaging illustrates its four-layered wall (Fig. 2D). The inner diameter of the tube is determined by the outer diameter of the mandrel, so the tube diameter can be facilely controlled just by using mandrels with proper diameters (Fig. 2A). PLGA 50 : 50 ES film tubes with the inner diameters of 10, 6, 4, and 1 mm and four layered walls are shown in Fig. 2E. It is also noted that in the whole fabrication process, we do not need expensive equipment or complex manipulations, and all operations are performed by hands under ordinary cell culture conditions. The whole manipulation time is quite short, normally less than 10 minutes for a tube. After we soaked these tubes in PBS and DMEM at 37 °C for 3 days, the scaffolds maintained the tubular shapes without layers detached (Fig. S7†). When we unroll the scaffolds, we could feel the strong adhesive force induced from the fibrin glue.

Fig. 2 Cell-free tubes. (A) The fabrication process. (B) The side view of a PLGA 50 : 50 (the mass ratio of PLA and PGA is 50 : 50) ES tube. (C) The cross section of the PLGA 50 : 50 ES tube. (D) The SEM image of the PLGA 50 : 50 ES tube. (E) PLGA 50 : 50 ES tube with different diameters, from left to right: 10, 6, 4, 1 mm. (F) Tubes of different materials with 3 mm diameter, from left to right: bacterial cellulose dried by air, PLGA 75 : 25 (the mass ratio of PLA and PGA is 75 : 25) ES, PLGA 75 : 25 cast, PCL ES. (G) PLGA cast tube with microwells. (H) PLGA cast tube with microgrooves.

To demonstrate the universality of our approach, we tried materials with different composition, fabrication method, and substructures. In some cases, microstructures are introduced into these materials to provide topological cues or facilitate nutrient transportation. For example, bacterial cellulose (BC), PLGA 75 : 25 ES film, PLGA 75 : 25 cast film, and polycaprolactone (PCL) ES film with flat surfaces were used for fabricating tubes with four layered walls (Fig. 2F). Also, PLGA cast and electrospun film with micro-scale structures were employed (Fig. 2G and H and 3D). We can obtain tubes with two layered walls containing regularly aligned pillars (Fig. 2G and S8†), grooves (Fig. 2H and S9†), and lattices (Fig. 3). This method provides a new strategy to rapidly fabricate multi-layered cell-free tubular structure with multiple adjustable parameters.

Fig. 3 Fabrication of tubes with lattices. (A) The fabrication process. (B) Optical imaging of the substrate of the ES film with lattice (the iron mesh net). (C) Optical imaging of the PLGA 50 : 50 electrospinning film with lattices on the substrate. (D) The rolled-up tube.

3.3. Fabrication of cell-containing scaffolds

Further, we can pattern blood vessel-related cells on the film to form blood vessel-like structures. The whole process includes two main parts: cell patterning and film rolling (Fig. 4A). The cell patterning part is performed on the flat film. Because techniques for cell patterning on 2D surfaces have been intensively developed recently, the whole process is convenient.^20,21 The PDMS chamber with individually addressable microchannels are placed on the flat film (which has been pretreated by fibronectin, FN), and three kinds of blood vessel-related cell lines: human umbilical vein endothelial cells (HUVECs), human aortic smooth muscle cells (HASMCs), and human aortic fibroblasts (HAFs), stained with different colours, are injected into the channels. The channels in the chamber are pre-designed such that the spacing along with the rolling direction equals to the perimeter of the mandrel, and the spacing perpendicular to the rolling direction equals to the length of the tube. After the cells attach onto the film (0.5–1 h after seeding), the PDMS chamber is peeled off, and the cells form confined geometrical patterns on the film (Fig. 4B). Soon afterwards, we can continue to perform the film rolling part. The glue is coated on both sides of the film with cells and the remaining rolling work is the same as previously described. The glue components and the resultant fibrin glue do not decrease the viability of the cells (Fig. S10†). After the formation of the tube, we fix and cut the tube into thin slices to take images. Confocal images show that the cells on different layers of the tube form well-organised structure, just as the cell distribution in the real blood vessel, where HUVECs, HASMCs, and HAFs are arranged from the innermost to outermost layers in the tube (Fig. 4C). All types of cells seeded maintain alive after a 3-d static culture in DMEM medium (Fig. S11†). Also, the same as the cell-free ones, the scaffolds sustain their tubular shape and integrity, and the glue keeps its stickiness (Fig. S7†). The whole process needs no more than 70 minutes, which is much faster compared to the reported methods.^{9,11–13,15,22–25} In clinics, it is very critical to obtain a scaffold with tailored size instantly. Our method exhibits a promising potential on this point. Cell density in each channel can be facilely, separately, and simultaneously regulated according to specific tissue organization. We can seed more ECs in the channel, because rapid and full coverage of endothelial cell layer is a key factor determining the patency of the blood vessel implants.^26,27 We can seed optimised amount of SMCs and FBs to best adapt to the growth of the engineered tissue.²⁸ The diameter, length, wall thickness, and mechanical strength of the tube can be easily controlled by material thickness and the size of the film. This method well balances the mechanical strength, structural similarity, as well as the fabrication time, and may become a novel design of tissue engineered blood vessel.

Fig. 4 Cell-laden tubes. (A) The fabrication process. (B) The film patterned with three kinds of vascular cells with three colours: red-HUVECs, green-HASMCs, and blue-HAFs. Insets: enlarged image of HUVECs, HASMCs, and HAFs. (C) The film after rolling with layered cells. Inset: enlarged view of the tube wall.

3.4. Reinforcement of other tubular scaffolds

As an extension of the fabrication, our strategy also allows a straightforward yet effective method to strengthen tubular scaffolds (Fig. 5A). For any scaffolds to be reinforced, we can selectively strengthen the tube on either its two shoulders or the whole tube by a sheath, according to our requirements (Fig. 5A). Here, we used a double-layered scaffold made of PLGA 50 : 50 ES film (thickness: 35 ± 2.6 μm, mean ± sd, n = 10) by above-mentioned protocols as the tube to be reinforced. We reinforce it using another two layers of PLGA 75 : 25 ES film (thickness: 36 ± 3.1 μm, mean ± sd, n = 10) by rolling the reinforcing materials around different sites of the tube. In order to realise the reinforcing process, we should insert a proper mandrel (its outer diameter should a little bit smaller than the inner diameter of the tube to facilitate the following protocol) into the tube to support the tube at first and the following steps are similar to those of tube preparation. For the sheath strengthening, we first prepare a piece of PLGA 75 : 25 ES film that can just be rolled around the tube with 2 layers. The two components of fibrin glue are coated on both sides of PLGA 75 : 25 film and the surface of the tube. Later, the film is rolled and glued on the whole outside wall of the tube. Finally, the supporting mandrel is pulled out. For the shoulder strengthening, the only differences are that two thin PLGA 75 : 25 stripes are used and that the outside wall at two extremes of the tube is reinforced. The enhancement of mechanical strengths: the suture retention and the burst pressure, two important mechanical factors for blood vessel implants, is obvious. Suture retention reflects the capacity that a scaffold holds the suture in surgical operation; burst pressure reflects the capacity that a scaffold withstands the pressure of liquid flow. After shoulder strengthening, the suture retention can increase from 0.31 ± 0.059 N (n = 3) to 1.04 ± 0.131 N (n = 3), an increase of almost 3.5 times. By contrast, the burst strength keeps almost unchanged (Fig. 5B and C). After sheath strengthening, the suture retention increases from 0.31 ± 0.059 N (n = 3) to 1.18 ± 0.194 N (n = 3), increasing almost 4 times higher, and meanwhile, the burst pressure increases from 0.0400 ± 0.01058 MPa (n = 3) to 0.141 ± 0.00814 MPa (n = 3), increasing to more than 3.5 times higher (Fig. 5B and C). From these data, we can see that strengthening on the shoulders will improve the suture retention only, while the strengthening by the sheath will enhance both the suture retention and burst pressure. This control offers us flexibility in scaffold reinforcement. Importantly, the whole strengthening process needs only around 10 minutes, and can be finished on site without disturbing the tube structures. Also, the choice of strengthening materials and the number of layers are quite flexible and the operation does not use any cell-unfriendly chemicals. This will provide us a general tool to selectively reinforce the specific part of the tissue engineering scaffolds on demand without impacting the structure and biocompatibility of the tube to be reinforced.

Fig. 5 Strengthened tubes. (A) The illustration of strengthening methods. U: unstrengthened tube, S1: strengthened on shoulders, S2: strengthened by sheath. (B) Burst pressure comparison of U, S1, and S2. (C) Suture retention comparison of U, S1, and S2.

3.5. Discussion

Development of a rapid and facile strategy to fabricate blood vessel-like structure has been a long-lasting goal in tissue engineering and regenerative medicine. Our strategy achieves an improved version by incorporating the advantages and learning the experiences of several previous studies in this aim.

First, in our strategy, fibrous biodegradable ES films are used as the major substrate material. These materials are easy to be produced in large scale, stored for a long time, and put to use readily. This will overcome the drawbacks using substrate from the biologically derived films, such as the cell-derived extracellular matrix (ECM), the SIS, and the amnion. For example, we have known that the production of SMC- or FB-containing ECM aforementioned needs about one month.^{5,7,11,12,29–31} This will become a critical speed-limiting step for the rapid fabrication of the blood vessel-like structures. Some approaches using SIS and the amnion from tissues to fabricate layered tubes.^32–34 Decellularization is involved in these approaches, which is complicated and time consuming. The post-treatment after the acquisition of the films needs around 3 days with multiple steps. On the other hand, it is known that the fibrous PLGA ES structure mimics the structure of ECM, which will facilitate the biomedical application, especially tissue engineering, of the scaffolds.^35,36 Because the interconnectedly porous structure of the ES film, the cells can communicate freely with each other through bioactive molecules they secreted. However, a problem rises that cells in different layers do not directly contact with each other because the barrier of the films. This issue may be solved in future explorations, because it is our hypothesis that the cells will form correct interactions as the degradation of the films and finally it will form an entire biological scaffold after the complete degradation of the polymers.

Second, the fibrin medical glue is used to accelerate the bonding of different layers and maintain the integrity of the whole structure. This glue provides us an effective method to hold the films in place within minutes. This saves much time compared with those strategies relying on cell–cell self-assembly. For example, the tissue fusion process of cell-containing ECM scaffolds needs around one month.^{5,7,11,12,29–31} This long period becomes another speed-limiting step in this strategy. For some strategies, the bonding of different layers just relies on compressing the tissue layers together. This may result in the isolation of the layers under physiological conditions with blood pressure, which is not allowed in the vascular graft. For example, in a human amnion-based rolling strategy, the decellularized amnion membrane was only rolled up around a steel mandrel without any fixation.^32,37 It has been reported that the fibrin glue can promote the tissue remodeling.^38–40 Therefore, using fibrin glue as the bonding reagent may be helpful for the subsequent implantation of the scaffolds. Moreover, because there is a thin liquid film of the glue between the mandrel and the cell surface before the drying of the glue, the cells can be protected from being crushed.

Third, cell suspension rather than cell sheets is used as the cell source. These cells will just need to be injected on the film. This process just needs 1–2 hours with minimum operations. Compared with rolling strategies based on cell sheets, our strategy improves the fabrication speed significantly owing to their lengthy cell sheets preparation time. The maturation of thick cell-containing ECM sheets needs around 4 weeks;^{5,7,11,12,29–31,41} the cell sheets peeled off from thermoresponsive culture dishes need around 4 days to get ready.^13,23,42 That means it is impossible to complete the tube fabrication within several hours. The cell sheet acquisition time becomes a speed-limiting step in their fabrication. This is the reason why some systems need much longer time to fabricate a tube compared with ours (several days vs. 70 min), though they have realised automation and applied similar bonding chemistry.²⁵

Fourth, PDMS stamps with well designed geometry can make the cell patterning of the film very precise. This is also partially attributed to the fact that the diameter of the tube in our system can be precisely controlled by the mandrel. This is a significant experience from our previous SIRM rolling strategy.² The well confinement of cells in different channels ensures the cell deposition in right position. Complex and time-consuming manipulations, such as rotating culture of ECs in the tube, are also omitted.^11,43 The adequate substrate elasticity avoids the hard-to-control shrinkage in SIMR rolling.^2,3,10 Also, the instant rolling after cell adhesion avoids the cell migration without control in a cell-containing hydrogel-ES membrane hybrid rolling strategy.¹³ Meanwhile, three kinds of vascular cells can be seeded all at once in our strategy (Fig. 4). This will not only save seeding time, but also benefit for the cell interactions with all their secreted factors communicating at the same time. In the well separated chamber, three kinds of cells can be seeded according to the optimal density in future application. For example, the critical factor for the success of artificial blood vessels, endothelialization, can be rapidly realised by increasing the number of ECs in corresponding channels (Fig. 4). Furthermore, using the chambers, the 2D film can be easily modified by cell matrix materials, growth factors, and/or small molecules before rolling, which provides us many choices to regulate the cell growth or remodelling behavior of the scaffolds.^44–49

Fifth, in our system, no equipment and special processing conditions are necessary. In our method, we can manually finish engineering of the tissue just relying on a mandrel. The only thing we need to pay attention to is each kind of cell in right amount and position. The fabrication condition is just the common cell culture environment. No special equipment is need. For example, compared with some automated strategies, no cell sheet rolling machine needs to be designed and fabricated.^14,25,50 Compared with 3D printing, the need of complex 3D printing equipment and expensive materials is eliminated.^51–54 This merit makes our strategy very practical and easy to learn, especially for those who have no extensive training and advanced working conditions.

Besides the items that need further tested above, our system has several points to be improved or certified. Firstly, in our system, the use of biomedical glue, FDA-approved polymer film, and the fabrication process without toxic components determine that the resultant blood vessel-like structure has good biocompatibility in principle. However, it should be certified in future animal experiments. Secondly, all the fabrication is completed manually, and thus the rolling speed and the force applied may be not always the same. Compared with those using rolling machine, the standardization of our production needs to be improved.^25,50 However, the standardization of the fabrication may comprise the simplicity of our method. Thirdly, it is not easy to realize differently oriented cells in our method. Normally, HUVECs are oriented longitudinally because the flow of the blood, while, SMCs are oriented circumferentially to support the vessel shrinkage and relaxation.¹ Incubating the scaffolds in a perfusion bioreactor would be helpful. The improved strategies are being conducted. For the present strategy, future work will focus on assessment of the mechanical properties of the scaffolds, cell behaviors in the scaffolds under culture, and performance of small animal implantation.

4. Conclusion

In conclusion, we have developed a rapid and facile strategy to fabricate tubular structures with controllably layered walls using biodegradable thin films, biomedical glue, and biocompatible smooth mandrel. We can develop fully biodegradable cell-free tubular structures within 10 minutes. The materials, diameter, wall thickness, and micropatterns of the tube can be freely selected. It will facilitate our preparation of adequate tubular structure according to different conditions. We can also obtain cell-laden tubular structures with multiple types of cells well-arranged in different layers within 70 minutes, which mimics the structure of tubular tissues in human. It provides a promising tool to create tailored tubular tissue with designed size within hours. We can also selectively strengthen any scaffolds using the rolling process. By strengthening on the sheath or shoulders of a tissue engineered scaffold, we can improve their suture retention and/or burst pressure on demand without impacting the structure and biocompatibility of the tube to be reinforced. Our method does not rely on any complex equipment and manipulations to realise the rolling process, which is cost-effective, easy to learn, and convenient to spread and use. We believe this novel tool will be generally useful in fast preparing and strengthening of tubular scaffolds.

Acknowledgements

This work was supported by the National Natural Science Foundation of China (31170905, 51373043, 81361140345, 21535001), the Chinese Academy of Science (XDA09030305), and the CAS/SAFEA International Partnership Program for Creative Research Teams.

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Footnote

† Electronic supplementary information (ESI) available. See DOI: 10.1039/c6ra12768a

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