VESSELS FOR SMALL DIAMETER APPLICATIONS

This invention concerns a composition consisting of biologically inert anionic polymers (alginate and carboxymethyl chitosan) and biologically active polymers (polyanionic inorganic polyphosphate and/or silica, and polycations, such as polyamines and/or polymers of basic amino acids, and/or synthetic oligo- or polypeptides containing the RGD motif, either alone or in combination) as a well as cross-linking calcium ions, that can be used for the fabrication of biologically active artificial blood vessels in particular in the small-diameter range. In addition, this invention concerns a device for the formation (pressing out) of the artificial blood vessels with the inventive composition. The novel blood vessels can be used as prosthetic vascular grafts in the surgical treatment of cardiovascular patients like patients requiring coronary and peripheral vessel reconstruction or aneurysm repair.

Background of the Invention

Cardiovascular diseases are the leading cause of mortality in the world. The number of patients both in Europe and in USA receiving prosthetic grafts for coronary and peripheral vessel reconstruction, aneurysm repair, or hemodialysis access is steadily increasing and expected to reach over 1.8 million today. However, until recently only limited clinical success for prosthetic vascular grafts has been achieved, in particular in the small-diameter range. The annual costs for those grafts increase to over 25 billion€ in Europe. Prosthetic vascular grafts are primarily sold for the application as endo vascular stent-grafts targeted to aortic aneurysm repair, peripheral vascular grafts, in hemodialysis access grafts and as small-diameter coronary artery bypass grafts. Since many patients, due to comorbid conditions or previous interventions, lack autologous vessels suitable for a surgical transfer to reconstruct/substitute damaged vascular segments an urgent clinical need for prosthetic substitutes that can compete with autologous vessels exists.

Vascular grafts

A suitable vascular graft should be readily available, being durable during long-time implantation, not eliciting an inflammatory potential, and not promoting thrombosis and/or infection. In addition, the graft wall should have similar mechanical properties like the native host vessel. At present, the most frequently applied materials for synthetic vascular grafts are made of expanded polytetrafluroethylene [ePTFE] and of polyethylene terephthalate [PET] (Bouten CVC, Dankers PYW, Driessen-Mol A, Pedron S, Brizard AMA, Baaijens FPT. Substrates for cardiovascular tissue engineering. Adv Drug Deliv Rev 2011;63:221-241). Both materials show sufficient long-term results for large-scale arterial reconstruction and large-diameter vessels (> 6 mm) but have been proven to display inferior performance and biological properties in small-diameter applications, e.g. coronary artery bypass grafting and arteriovenous shunts (Hasan A, Memic A, Annabi N, Hossain M, Paul A, Dokmeci MR, Dehghani F, Khademhosseini A. Electrospun scaffolds for tissue engineering of vascular grafts. Acta Biomater 2014;10: 11-25). The failure in the latter applications of these synthetic implants are caused by an increased surface thrombogenicity, due to the lack of a functional endothelium, and a substantial development of intimal hyperplasia as a consequence of chronic inflammations (Klinkert P, Post PN, Breslau PJ, van Bockel JH. Saphenous vein versus PTFE for above-knee femoropopliteal bypass. A review of the literature. Eur J Vase Endovasc Surg 2004;27:357-362; Loh SA, Howell BS, Rockman CB, Cayne NS, Adelman MA, Gulkarov I, Veith FJ, Maldonado TS. Mid- and long-term results of the treatment of infrainguinal arterial occlusive disease with precuffed expanded polytetrafluoroethylene grafts compared with vein grafts. Ann Vase Surg 2013;27:208-217). A severe consequence of the lack of a functional endothelial layer is an increased microbial contamination, which often needs implant replacement.

At present small diameter vascular grafts with an internal diameter of less than 5 mm are not approved by the FDA for clinical use due to their high failure rates (Linderman S, Araya J, Pathan S, Nelson D, Phaneuf M and Contreras M. A small diameter bioactive prosthetic vascular graft with activated protein C (546.9). FASEB J 2014;28: Suppl 546.9). The major reason is that those grafts have the potential to promote thrombosis (Enomoto S, Sumi M, Kajimoto K, Nakazawa Y, Takahashi R, Takabayashi C, Asakura T, Sata M. Long-term patency of small-diameter vascular graft made from fibroin, a silk-based biodegradable material. J Vase Surg 2010;51 : 155-164) as well as to enhance endothelial cell proliferation (He W, Hu Z, Xu A, Liu R, Yin H, Wang J, Wang S. The preparation and performance of a new polyurethane vascular prosthesis. Cell Biochem Biophys 2013;66:855-866). The invention described here concerns a novel type of small diameter vascular grafts, termed biomimetic tissue-engineered blood vessels [bTEBV], with different bioactive functionalities. The bTEBV meet the needs of individual patients. They are fabricated in a modular way: a) a universal, inert scaffold is used, into which

b) bio functionally active polymers have been incorporated.

a) Universal inert scaffold: The hydrogel backbone of the bTEBV is built by two natural polymers: (i) alginate, a natural plant polymer, and (ii) a modified chitosan that is produced, after deacetylation, from chitin, the second most abundant biopolymer. The OH groups and/or N¾ groups initially present in chitosan had been converted to carboxymethyl groups, allowing the binding to cations. The resulting modified chitosan is termed carboxymethyl chitosan [CMC]; the use of N, O-carboxymethyl chitosan [NO-CMC] has been described in the Examples.

This biomaterial, consisting of alginate and NO-CMC, can be hardened via formation of Ca ²⁺ bridges to a durable hydrogel. The mechanical properties of the resulting material (alginate- Ca ²⁺ -NO-CMC) can be tuned and easily adjusted. The material is durable. As described in the Examples, the vascular grafts can be functionally employed with a pulsatile, rhythmic flow for over 4 weeks without leakage. b) Embedded bio functionally active polymers: The basic alginate-Ca ²⁺ -N O-CMC scaffold is biologically inert. As a first component with biological activity gelatin is added. Gelatin exposes the Arg-Gly-Asp [RGD] cell recognition signal and, by that, allows the cell- membrane bound integrins to bind.

As the second biologically active component inorganic polyphosphate [polyP] is added that is exposing negatively charged groups, like alginate and NO-CMC. PolyP has been described as a key modulator in platelet-mediated pro-inflammatory and pro-coagulant disorders (Muller F, Mutch NJ, Schenk WA, Smith SA, Esterl L, Spronk HM, Schmidbauer S, Gahl WA, Morrissey JH, Renne T. Platelet polyphosphates are proinflammatory and procoagulant mediators in vivo. Cell 2009;139:1143-1156).

WO 03/080142 discloses a hybrid artificial blood vessel including a biodegradable polymer supporting layer and the manufacturing process thereof. The inventors co-added polyP (as Na -polyP) to the alginate and N.O-CMC during the preparation. This scaffold backbone of the bTEBV, consisting of the above mentioned polymers, all of them are polyanionic, can be hardened by cationic Ca ²⁺ .

One aspect of this invention is the unexpected finding that this basic material, if prepared in a fixed sequence of steps, as outlined in the Examples (see also Figure 6), can be used for fabrication of artificial blood vessels, by pressing out the material through a newly developed extruder, being part of this invention. This property of the material was surprising because in patent application GB1413154.4 (Printable morphogenetic phase-specific chitosan-calcium- polyphosphate scaffold for bone repair; inventors: Miiller WEG, Schroder HC, Wang XH), it has been reported that this material [N, O-CMC-Ca ²⁺ -polyP] causes an increased mineralization of cells, an effect that would be contradictory to its use as a flexible material for prosthetic vascular grafts for long-term treatment of, for example, atherosclerotic patients.

Another aspect of this invention is the unexpected finding that the stimulatory effect on blood coagulation of the material, which is of advantage for application of this material in bone tissue engineering and repair but of extreme disadvantage for application in blood vessel reconstruction (risk of thrombosis), can be abolished by addition of polyamines, polycationic amino acids or a His/Gly-tagged RGD peptide.

Polycationic amino acids, like poly(L-Lys) and poly(D-Lys), support the attachment of cells to natural and artificial surfaces followed by an activation of cell metabolism. The inventors also incorporated into the bTEBV scaffold the peptide NH ₂ - RGDGGRGDGGGRGDGGGHHHHHH-COOH (SEQ ID No. 1) [termed His/Gly-tagged RGD] which contains three copies of the Arg-Gly-Asp [RGD] cell recognition signal, in order to increase the propensity of the endothelial cells to adhere to the artificial vessel surface. The amino acid His has been included into the peptide in order to utilize the property of this amino acid to form weak interactions with alginate (Tielen P, Kuhn H, Rosenau F, Jaeger KE, Flemming HC, Wingender J. Interaction between extracellular lipase LipA and the polysaccharide alginate of Pseudomonas aeruginosa. BMC Microbiol 2013;13: 159. doi: 10.1186/1471-2180-13-159) and to bind to negatively charged bio-polymer surfaces (MacQuarrie JL, Stafford AR, Yau JW, Leslie BA, Vu TT, Fredenburgh JC, Weitz JI. Histidine-rich glycoprotein binds factor Xlla with high affinity and inhibits contact-initiated coagulation. Blood 2011;117:4134-4141). Surprisingly, it was now found that the biological activities of the polyanion polyP and of the polycations poly(L-Lys) and poly(D-Lys), and of the His/Gly-tagged RGD peptide, present in the inventive material used for the fabrication of the bTEBV, are not abolished in the mixture of the polyanionic (polyP, alginate, N,O-CMC polymer) and polycationic components [poly(L-Lys) and poly(D-Lys), His/Gly-tagged RGD peptide], as well as in the presence of calcium ions.

A scheme of the composition of the inventive bTEBV to be used as prosthetic vessel is given in Figure 1A; it outlines the functional arrangement of the components, composing the prosthetic vascular graft: scaffold backbone formed by the inert anionic polymers, N,O-CMC and alginate, enclosing the embedded bio functionally active polymers, anionic polyP and gelatin, that are linked together via the divalent cations. Incorporated into this scaffold are poly(L-Lys), poly(D-Lys) and/or the His/Gly-tagged RGD peptide, which promote the adhesion of endothelial cells.

Polyphosphate (polyP)

Previous studies revealed that PolyP is a linear polymer that can be synthesized both chemically and enzymatically (reviewed in: Schroder HC, Miiller WEG, eds. Inorganic Polyphosphates - Biochemistry, Biology, Biotechnology. Prog Mol Subcell Biol 1999;23:45- 81).

The enzymatic formation and degradation of polyP is catalyzed by polyphosphate kinases and exo- and endopolyphosphatases (for example: Lorenz B, Miiller WEG, Kulaev IS, Schroder HC. Purification and characterization of an exopolyphosphatase activity from Saccharomyces cerevisiae. J Biol Chem 1994;269:22198-22204).

The alkaline phosphatase is an exopolyphosphatase that degrades polyP to monomeric phosphate (Lorenz B, Schroder HC. Mammalian intestinal alkaline phosphatase acts as highly active exopolyphosphatase. Biochim Biophys Acta 2001;1547:254-261).

PolyP is present in animal tissue (Leyhausen G, Lorenz B, Zhu H, Geurtsen W, Bohnensack R, Miiller WEG, Schroder HC. Inorganic polyphosphate in human osteoblast-like cells. J Bone Mineral Res 1998;13:803-812; Schroder HC, Kurz L, Miiller WEG, Lorenz B. Polyphosphate in bone. Biochemistry (Moscow) 2000;65:296-303) and in platelets (Smith SA, Mutch NJ, Baskar D, Rohloff P, Docampo R, Morrissey JH. Polyphosphate modulates blood coagulation and fibrinolysis. Proc Natl Acad Sci USA 2006;103:903-908).

PolyP is morphogenetically active after complex formation with Ca ²⁺ ions (polyP ^» Ca ²⁺ - complex or polyP ^» Ca ²⁺ -salt) and displays the following activities:

- induction of expression and stimulation of activity of the alkaline phosphatase (Muller WEG, Wang XH, Diehl-Seifert B, Kropf K, SchloBmacher U, Lieberwirth I, Glasser G, Wiens M, Schroder HC. Inorganic polymeric phosphate/polyphosphate is an inducer of alkaline phosphatase and a modulator of intracellular Ca ²⁺ level in osteoblasts (SaOS-2 cells) in vitro. Acta Biomater 2011;7:2661-2671);

- stimulation of expression of BMP-2 (Wang XH, Schroder HC, Diehl-Seifert B, Kropf K, SchloBmacher U, Wiens M, Muller WEG. Dual effect of inorganic polymeric phosphate/polyphosphate on osteoblasts and osteoclasts in vitro. J Tissue Eng Regen Med 2013;7: 767-776);

- induction of hydroxy apatite formation (Wang XH, Schroder HC, Diehl-Seifert B, Kropf K, SchloBmacher U, Wiens M, Muller WEG. Dual effect of inorganic polymeric phosphate/polyphosphate on osteoblasts and osteoclasts in vitro. J Tissue Eng Regen Med 2013;7: 767-776).

The following patent publications appear to be relevant: GB1406840.7 Morphogenetically active hydrogel for bioprinting of bio artificial tissue. Inventors: Muller WEG, Schroder HC, Wang XH; GB 1403899.6 Synergistic composition comprising quercetin and polyphosphate for treatment of bone disorders. Inventors: Muller WEG, Schroder HC, Wang XH; and GB1319416.2. Modulator of bone mineralization based on a combination of polyphosphate/carbonate and carbonic anhydrase activators. Inventors: Muller WEG, Schroder HC, Wang XH.

Biosilica

Biosilica is an enzymatically formed natural polymer that is synthesized by the enzyme silicatein (Muller WEG, Schroder HC, Burghard Z, Pisignano D, Wang XH. Silicateins: a paradigm shift in bioinorganic chemistry. Enzymatic synthesis of inorganic polymeric silica. Chem Eur J 2013;19:5790-5804). Biosilica induces the expression of cytokines / growth factors in bone-forming cells (BMP-2 and osteoprotegerin) and inhibits the differentiation of pre-osteoclasts into mature osteoclasts (reviewed in: Wang XH, Schroder HC, Wiens M, Ushijima H, Muller WEG. Bio-silica and bio-polyphosphate: applications in biomedicine (bone formation). Curr Opin Biotechnol 2012;23:570-578).

The following patents and patent applications appear to be relevant: EP1320624; US7169589B2 Silicatein-mediated synthesis of amorphous silicates and siloxanes and their uses. Inventors: Muller WEG, Lorenz A, Krasko A, Schroder HC; US6670438B1 Methods, compositions, and biomimetic catalysts for in vitro synthesis of silica, polysilsequioxane, polysiloxane, and polymetallo-oxanes. Inventors: Morse DE, Stucky GD, Deming, TD, Cha J, Shimizu K, Zhou Y; DEI 0246186 In vitro and in vivo degradation or synthesis of silicon dioxide and silicones, useful e.g. for treating silicosis or to prepare prosthetic materials, using a new silicase enzyme. Inventors: Muller WEG, Krasko A, Schroder HC; EP1546319 Abbau und Modifizierung von Silicaten und Siliconen durch Silicase und Verwendung des reversiblen Enzyms. Inventors: Muller WEG, Krasko A, Schroder HC; EP1740707 Enzym- und Template-gesteuerte Synthese von Silica aus nicht-organischen Siliciumverbindungen sowie Aminosilanen und Silazanen und Verwendung. Inventors: Schwertner H, Muller WEG, Schroder HC; EP09005849.6 Use of silintaphin for the structure-directed fabrication of (nano)composite materials in medicine and (nano)technology. Inventors: Wiens M, Muller WEG, Schroder HC, Wang X; DE102004021229.5 Enzymatic method for producing bioactive, osteob last-stimulating surfaces and use thereof. Inventors: Muller WEG, Schwertner H, Schroder HC; US 60839601 Biosilica-adhesive protein nano-composite materials: synthesis and application in dentistry. Inventors: Muller WEG, Schroder HC, Geurtsen WK; WO 2010/036344 Compositions, oral care products and methods of making and using the same. Inventors: Miller J, Hofer H, Geurtsen W, Lucker P, Wiens M, Schroder HC, Muller WEG; GB 1405994.3 Osteogenic material to be used for treatment of bone defects. Inventors: Muller WEG, Schroder HC, Wang XH; and GB1408402.4 3D cell printing of bioglass-containing scaffolds by combination with cell-containing morphogenically active alginate/gelatin hydrogels. Inventors: Muller WEG, Schroder HC, Wang XH.

Detailed description of the invention

This invention relates to the production of a novel type of artificial small diameter blood vessels, termed biomimetic tissue-engineered blood vessels [bTEBV], with a modular composition, according to the appended claims. In particular, the inventive bTEBV are composed of:

a) a universal, inert hydrogel scaffold that consist of the basic polymeric anionic (negatively charged) nature-derived polymers, alginate and modified chitosan (carboxymethyl chitosan), b) the biologically active polymers, inorganic polyP (polyanionic - negatively charged), and/or silica and gelatine, and

c) calcium ions.

In addition, polycations, such as polyamines or polymers of basic amino acids, or synthetic oligo- or polypeptides containing the RGD motif, either alone or in combination, are added.

The carboxymethyl chitosan can be formed by carboxymethylation of the amino groups of chitosan (N-carboxymethyl chitosan) or the hydroxy groups of chitosan (O-carboxymethyl chitosan) or both (N, 0-carboxymethyl chitosan). In the Examples as given below, bTEBV containing N, 0-carboxymethyl chitosan (N, O-CMC) are described.

The non-carboxymethylated amino groups and / or the non-carboxymethylated hydroxy groups of the carboxymethyl chitosan can be acetylated or partially acetylated.

The polymeric silica or biosilica can be formed enzymatically by silicatein. Gelatin, embedded as the third bio functionally active biopolymer to the biologically inert alginate - CMC scaffold, besides polyP and silica, exposes the cell recognition signal, Arg-Gly-Asp [RGD]. Instead of gelatine, also other collagen-derived products can be used. The average chain lengths of the polyP molecules can be in the range 10 to up to 100 phosphate units. PolyP molecules with an average chain length of about 40 phosphate units are preferably used.

The biosilica can be formed by an enzyme involved in biosilica metabolism, such as silicatein or a silicatein fusion protein. The silicatein polypeptide or a silicatein fusion protein can be produced by using a prokaryotic or an eukaryotic expression system, or can be produced synthetically fo lowing established procedures (see above-mentioned patents / patent applications on silicatein). The inventors unexpectedly found that this material can be used for the fabrication of biologically active vascular grafts, which are superior to conventional vascular grafts used for example in bypass surgery, by applying a special extruder, being part of this invention, as described below. Thereby the basic material is hardened by exposure to Ca ²⁺ through formation of Ca ²⁺ bridges between the polyanions, alginate, Ν,Ο-CMC, and polyP. The bTEBV are formed by pressing the hydrogel through the extruder into a hardening solution, containing the cation Ca ²⁺ . During this process the polymers turn from a randomly oriented pattern to a concerted, oriented arrangement (see, e.g., Figure 10). The scaffold becomes solid, durable and biologically active.

In this universal scaffold [alginate-Ca ²⁺ -N,O-CMC-polyP-silica], polycations such as poly(L- Lys), poly(D-Lys) or a His/Gly-tagged RGD peptide (three RGD units) are incorporated, which promote the adhesion of endothelial cells to the bTEBV surface.

Instead of basic polyamino acids [poly(L-Lys), poly(D-Lys), poly(L-Arg), poly(D-Ag)], other polycations such as the polyamines, e.g., spermine, spermidine, cadaverine or putrescine, can be incorporated in the universal scaffold, or other RGD peptides than that described in the Examples can be used equally well.

The analysis of the mechanical properties of the universal scaffold material underlying the bTEBV (alginate-Ca ²⁺ -N,0-CMC-polyP-silica) revealed a hardness (elastic modulus) of 475 kPa even after a short (5 min) incubation in the CaCl ₂ solution that can be further increased by a extension of the incubation period or addition of ethanol.

The inventive material can be used in particular for the fabrication of artificial blood vessels (bTEBVs) in the small diameter range. As described under Examples, bTEBVs with an outer size 6 mm and 1.8 mm, and an inner diameter 4 mm and 0.8 mm, respectively, are durable in 4-week pulsatile flow experiments at an alternating pressure between 25 and 100 mbar (18.7 and 75.0 mm Hg). The burst pressure of the larger (smaller) vessels is 850 mbar (145 mbar).

The biological activity of the scaffold has been demonstrated in cell growth studies, as described under Examples. The cell adhesion behaviour of this matrix is suitable for endothelial cells, as demonstrated for human, umbilical vein/vascular endothelial cells, HUVEC (Figures 1B-D; see also Figures 10 and 11). Incorporation of polycationic poly(L-Lys), poly(D-Lys), and especially the His/Gly-tagged RGD peptide, markedly increases the adhesion of HUVEC cells to the surface of the bTEBV hydrogel matrix.

Based on these properties the metabolically degradable polymeric backbone building the bTEBV is a suitable material for prosthetic vascular grafts.

The inventive device for the fabrication of the artificial vascular grafts (bTEBV) composed of the inventive material, the extruder, comprises or consists of the following parts (for a photo and a technical drawing, see Figures 3 and 4):

- a syringe containing the CMC-based material, which is hooked to the extruder,

- a plunger for pressing the material into the syringe and through the aperture disc,

- an aperture disc (see Figure 5), hooked within a molder,

- the molder,

- the hardening tube, into which the inventive material is squeezed via the aperture disc, and

- a central stab sitting within the central guide of the aperture disc (see Figure 5) and extended into the hardening tube, around which the vessel is formed.

In a previous patent application (GB1413154.4; Printable morphogenetic phase-specific chitosan-calcium-polyphosphate scaffold for bone repair. Inventors: Miiller WEG, Schroder HC, Wang XH), the inventors disclosed a formula for the synthesis of a biocompatible, biodegradable, and printable material, composed of alginate, CMC and Na-polyP; the hardening of this material by exposure to calcium ions; thereby the Na ⁺ cations in the polyP are exchanged by Ca ²⁺ allowing the bridging of polyP to CMC and rendering the composite material particularly stable without loosing the biological activity of polyP; and the biological function of the CMC-polyP printed material to induce bone cells to biomineralization, and to accelerate the clotting process of human blood. However, the material as described in GB1413154.4, markedly enhanced blood coagulation, an effect which is desired for application in bone tissue engineering and repair (addressed in GB1413154.4) but is regarded as absolutely detrimental for application in blood vessel reconstruction (as in the present patent application). PolyP-Ca ²⁺ -N,0-CMC, one component of the inventive material, involves the risk to induce thrombosis in patients after blood vessel implantation.

However, unexpectedly, the material according to this invention, used for the formation of the bTEBV, does not induce blood coagulation. The finding that the pro-thrombotic effect of polyP is abolished in the presence of polyamines or polycationic amino acids [for example: poly(L-Lys) and poly(D-Lys)] or the described His/Gly-tagged RGD peptide was surprising because it could not be expected that the polycations present in the inventive material will behave differently compared with the cation Ca ²⁺ with regards to the effect on blood coagulation. In addition, it was surprising that the biological activities of the polyanion polyP and of the polycations poly(L-Lys) and poly(D-Lys), and the His/Gly-tagged RGD peptide are not abolished in the mixture of both polyanionic (polyP, alginate, and N,O-CMC polymer) and polycationic components [poly(L-Lys) and poly(D-Lys), and His/Gly-tagged RGD peptide], as well as in the presence of calcium ions.

One striking advantage of the bTEBV according to this invention is that they can be produced over a wide range of internal and outer diameters, by varying the distance of the center of the central stab to the outer edge of the outlets of the aperture disc or by varying the diameter of the central stab (see Figure 5). In particular, small diameter vessels in the size range < 5 mm can be fabricated, a size range that has turned out to be critical with regard to potential thrombotic complications as reported for synthetic graft materials. In Figure 5 an aperture disc for fabrication of vessels with an outer diameter of 1.8 mm and an inner diameter of 0.8 mm is shown.

Another advantageous property of the inventive material, used for the formation of the bTEBV, is the fact that its mechanical properties such as hardness / elastic modulus can be tuned and adapted to the needs of individual patients and applications. For example, the hardness can be adjusted to values existing in arteries and veins, by changing the duration of the CaCl ₂ incubation; even more, the material can be additionally hardened by co-addition of ethanol. By that the bTEBV can be used for a personalized application as an anatomically adequate implant.

An advantage of the inventive bTEBV is that cationic polymers, e.g. poly(L-Lys), poly(D- Lys), and especially His/Gly-tagged RGD can be incorporated. These additives turn the material to a suitable template for the endothelial cells, since these cells readily attach on those derivatized matrices. This functionalization prevents the formation of thrombosis after (potential) grafting.

A further favorable aspect of the inventive novel type of artificial blood vessels is that the polymeric materials used for building these vessels are degradable (see Figure 2). It has been reported that degradation of alginate-based bio materials in vivo occurs after disintegration of the material in response to a gradual exchange of gelling Ca ²⁺ ions with Na ⁺ . Chitosan is metabolically hydrolyzed by macrophages via chitotriosidase and polyP is degraded by ALP (Leyhausen G, Lorenz B, Zhu H, Geurtsen W, Bohnensack R, Miiller WEG, Schroder HC. Inorganic polyphosphate in human osteoblast-like cells. J Bone Mineral Res 1998;13:803- 812). Finally, biosilica is assessable to the carbonic anhydrases, a family of enzymes that is related to the silicase (Schroder HC, Krasko A, Le Pennec G, Adell T, Hassanein H, Miiller IM, Miiller WEG. Silicase, an enzyme which degrades biogeneous amorphous silica: Contribution to the metabolism of silica deposition in the demosponge Suberites domuncula. In: Miiller WEG (2003) [ed] Silicon Biomineralization. Berlin: Springer Press; Progress Molec Subcell Biol 2003;33:249-268). The removal of Ca ²⁺ from the "stabile" polymer complexes within the universal scaffold might be achieved by oxalate chelation as well by organic polymers in the extracellular environment.

The inventive material as well as the inventive device used for the fabrication of tube-like structures using this material can also be used for further applications, including the formation of Schwann cell like structures promoting the ingrowth of nerve cell axons and supporting nerve regeneration. In the vertebrate peripheral nervous system (PNS), Schwann cells form the myelin sheath around the axons of the motor and sensory neurons. They are involved not only in the formation of a nerve extracellular matrix but also in the conduction of nervous impulses along the axons. The inventive tubes formed by the device according to this invention provide not only a "tunnel" guiding the growth of the axon toward its target (neurons, muscles, organs) but also insulation to the axon like the natural myelin sheath. The nodes of Ranvier, i.e. the gaps between adjacent Schwann cells that allow the saltatory conduction of the action potential can be produced by consecutive fabrication of the inventive tubes, e.g., by altering the composition (e.g., hardness, degradability) of the inventive material between each step (i.e., during pressing-out of the tubes) or mechanically after fabrication of the tubes (i.e., after pressing-out of the tubes). Potential applications comprise, among others, the use of the tubes as implants after spinal cord injury or in patients suffering from neuropathies involving Schwann cells, e.g. patients with multiple sclerosis.

Another example is the formation of flexible light transmitting structures for diagnostic, surgical and therapeutic applications, for example in endoscopy, in particular disposable endoscopy, e.g. arterial endoscopy (arterioscopy, cardiovascular endoscopy), pulmonary endoscopy (bronchoscopy), urinary endoscopy (cystoscopy), esophagogastroduodenoscopy (e.g., enteroscopy or intraoperative cholangioscopy), arthroscopy, otoscopy, and laparoscopy, as well as non-medical uses of endoscopy (e.g. borescopy). Potential therapeutic applications also include, in addition to nerve restoration (see above), a communication between the nervous system and the prosthesis in patients after amputations, or communication with body- implanted sensors or opto-electric pulse makers.

Light transmitting properties of the material, as well its refractive index, can be modulated by embedding silicatein (EP1320624; Silicatein-mediated synthesis of amorphous silicates and siloxanes and their uses. Inventors: Muller WEG, Lorenz A, Krasko A, Schroder HC) in the alginate-Ca ²⁺ -N, C-CMC-polyP matrix and incubation of the tubes, after pressing-out using the inventive device, with a suitable substrate, e.g. prehydrolyzed tetraethoxysilane (TEOS), in the absence or presence of dopants, if required, following established procedures (Muller WEG, Link T, Schroder HC, Korzhev M, Neufurth M, Brandt D, Wang XH. Dissection of the structure-forming from the structure-guiding activity of silicatein: a biomimetic molecular approach to print optical fibers. J Mater Chem B 2014;2:5368-5377; Muller WEG, Tolba E, Schroder HC, Diehl-Seifert B, Link T, Wang XH. Biosilica-loaded poly(8-caprolactone) nano fibers mats provide a morphogenetically active surface scaffold for the growth and mineralization of SaOS-2 cells. Biotechnol J 2014;9: 1312-1321).

The invention will now be described further in the following preferred examples, nevertheless, without being limited thereto. For the purposes of the present invention, all references as cited herein are incorporated by reference in their entireties. In the Figures, Figure 1 shows the production of biomimetic tissue-engineered blood vessels [bTEBV]. (A) The scaffold is composed of the negatively charged inert polymers, N, O-carboxymethyl chitosan (N. O-CMC) and alginate. This scaffold is enriched with the bioactive polyanionic polyP, as well as gelatin. The hydrogel is pressed through the inventive extruder and immediately submersed in a solution, containing the cation Ca ²⁺ , followed by hardening of the material. The Ca ²⁺ can be partially substituted by polycations, e.g. poly(L-Lys), or the peptide His/Gly-tagged RGD. After addition of poly(L-Lys), especially together with the His/Gly- tagged RGD peptide, the endothelial cells densely cover the surface of the hydrogel. The outer cell layer, composed of mural cells, is only sketched for completeness, but not studied here. (B to D) Human, umbilical vein/vascular endothelial cells, HUVEC cells, grown for 2 weeks on the following scaffold matrices; (B) basic scaffold alone, (C) basic scaffold, supplemented with poly(D-Lys), or (C) basic scaffold, supplemented with His/Gly-tagged RGD. The cells are stained, after fixation, with DRAQ5 (blue fluorescence) and labeled antibodies against actin (red).

Figure 2 shows the composition of the polyanionic scaffold material, N,O-CMC, alginate, polyP and silica in the basic scaffold of the bTEBV. Due to ionic linkages the cation Ca ²⁺ crosslinks those polymers and forms a more solid implant material. The durable hydrogel formed is proposed to be metabolically dissolved, via glycosidases and alkaline phosphatases (as well as carbonic anhydrases) to the respective monomers. The exchange of Ca ²⁺ by Na ⁺ is assumed to be mediated by both inorganic and organic chelators.

Figure 3 shows a technical drawing of the extruder used for the fabrication of the bTEBV. The N, O-CMC-based material is filled into a syringe (s) which is hooked to the inventive extruder (et). (A and B) By pressing the material with the plunger (p) into the syringe and through the aperture disc (ad), hooked within the molder (m) via the aperture disc in the molder, into the hardening tube, a tube around the central stab is formed. The tube, the bTEBV, undergoes hardening in the hardening tube, a process driven by Ca ²⁺ .

Figure 4 shows the experimental fabrication of the bTEBV in the inventive extruder. (A) The extruder device with its main parts, the plunger (p), syringe (s), molder (m) and the prepared bTEBV in the hardening tube (ht), is shown. (B, D, G) The ready fabricated bTEBV is pulled over a tube outlet of the pressure device (pd). In (G) the cell culture medium that has been pumped into the vessels is colored in purple. Close up of the (C) central part of the extruder, with the syringe (s) and the molder (m), and (E, F, H) the finally fabricated bTEBVs with an inner diameter of 0.8 mm are depicted.

Figure 5 shows the aperture disc, hooked within the molder, through which the inventive material is squeezed, with the central stab sitting within the central guide of the aperture disc (Left). The inventive material fuses after passing the outer holes of the aperture disc around the central stab and hardens in the presence of calcium ions within the hardening tube, under formation of a tube (vessels). (Right) The aperture disc for fabrication of vessels with an outer diameter of 1.8 mm and an inner diameter of 0.8 mm is shown.

Figure 6 shows the sequential addition of the polymers to prepare the hydrogel: (A and B) Suspending of N.O-CMC and silica in saline, and preparation of a solution; (C and D) addition of Na-polyP (polyP) and preparation of a homogeneous solution, again by stirring; (E and F) supplementation of solid sodium alginate (alg), together with low-melting gelatin (gel), resulting in a viscous gel. (G) Filling of the hydrogel, the scaffold to be used for fabrication of the bTEBV, in the extruder or for bioprinting, into a syringe. (H) Printed hydrogel discs, prior (left) or after (right) hardening with CaCl ₂

Figure 7 shows that changing the aperture disk within the extruder housing from the one with a 0.8 mm central stab resulting in a smaller material extrusion to aperture discs with a 4 mm central stab as well as equally enlarged material outlets allows the fabrication of bTEBV with an inner diameter of 4 mm (right vessel) instead of the 0.8 mm sized bTEBV (inner diameter). Left vessel: commercially available vessel grafts, totally prepared from non- biological polymers.

Figure 8 shows the pressure testing of the bTEBV. Here the small sized vessels, 1.8 mm [inner diameter 0.8 mm] were examined. (A) Device for studying the stability of the vessels. The two ends of the respective bTEBV are connected via hoses to the connecting tubes (c) that are in full close with the pump (p). The manometer (ma) is connected with the water flow by a Y connection (y). Colored (purple) medium is alternatively pumped into the vessels; it is temperature-controlled at 37°C, using a thermostat (t). (B to E) During the pumping cycle the outer diameter of the vessels varies between 1.80 mm and 2.05 mm.

Figure 9 shows the determination of the elastic modulus of the biomaterial, used for fabrication of the bTEBV. Hardening of the material is performed using 2.5% CaCl ₂ in 70% ethanol. The measurements are performed at 10°C (open bars), 30°C (closed) or 50°C (cross- hatched), using the ferrule-top nanoindenter. The determinations are performed in saline; the moduli are given in kPa. Significant values with respect to the values measured at 15 s and 30°C are marked; * P < 0.01. Figure 10 shows the influence of the different additives to the universal, basic scaffold, prepared from N,0-CMC, silica, polyP and alginate/gelatin in a fixed sequence, on cell growth/viability. The adhesion-promoting oligo/polymers, the polycationic poly-L-lysine and poly-D-lysine as well as His/Gly-tagged RGD were incorporated into the scaffold as described under Methods. The biomaterial was printed to 0.95 to 1.0 cm discs and placed into 48 well plates. After an incubation period of 7 d the growth/viability of the cells was determined using the XTT assay and the absorbance was measured at 490 nm. Both the "control-scaffold" (cont-sca) and the bioactive scaffolds "poly(L-Lys)-scaffold" (poly(L-Lys)-sca), "poly(D- Lys)-scaffold" (poly(D-Lys)-sca) and "RGD-scaffold" (RGD-sca) were examined. After an incubation period of 7 d the amount of insoluble tetrazolium salt formed was determined. The standard errors of the means (SEM) are indicated (n = 10 experiments); * P < 0.05.

Figure 11 shows the density of HUVEC cells grown on different scaffolds for 2 weeks. Left panel: staining with DRAQ5; right panel: staining for the cytoskeleton structures, actin; the fluorescent immunostained cells are shown. Density of the cells (A and B) onto "control- scaffold", (C and D) onto "poly(D-Lys)-scaffold", (E and F) onto "poly(L-Lys)-scaffold", (G and H) onto "RGD-scaffold" and finally (I and J) onto a scaffold containing both poly(L-Lys) and His/Gly-tagged RGD.

Examples

In the following examples, only a method is described with using N,0-CMC and polyP of an average chain of 40 phosphate units. Nevertheless, the method according to this invention can also be applied using O-carboxmethyl chitosan (O-CMC) and N-carboxymethyl chitosan (TV- CMC), and polyP molecules with chain lengths other than 40 phosphate units, and the person of skill will be able to adjust the method as described accordingly.

Preparation of the TV, O-carboxymethyl chitosan-based scaffold

The universal scaffold is prepared from NO-CMC, silica, polyP, alginate/gelatin in a fixed sequence, as outlined under Methods (Figure 6). The hydrogel is prepared from the biologically inert polyanionic polymers, N, O-carboxymethyl chitosan (NO-CMC) and silica, suspended in saline (Figure 6A and B). During stirring a homogeneous, viscous solution of these anionic polymers is obtained that is supplemented with the bioactive, anionic, polymers polyP, silica and gelatin, as outlined under Methods. The sequence of addition of the components is crucial (Figure 6). At first N.O-CMC and silica are suspended in saline and then stirred until homogeneity (Figure 6A and B). In the second step solid Na-polyP is added which can be brought to a homogeneous gel (Figure 6C and D). In the third step solid sodium alginate together with low-melting gelatin is added, resulting in a highly viscous gel (Figure 6E and F). This hydrogel is then filled into the syringe (Figure 6G) that is fitted with its plunger to the home-made extruder, to fabricate the vessel or in the bioprinter, to print the gel discs (Figure 6H), after centrifugation (removal of air bubbles).

The bio functionally active polymers can be embedded into the scaffold either during its preparation (His/Gly-tagged RGD) of after completion and hardening of the scaffold (poly(L- Lys) and poly(D-Lys)). Accordingly, three types of scaffolds are described in the Examples; "control scaffold" containing N,0-CMC, silica, polyP and gelatin, and scaffolds enriched with the active polymers poly(L-Lys), poly(D-Lys) and His/Gly-tagged RGD that are termed "poly(L-Lys)-scaffold", "poly(D-Lys)-scaffold", and "RGD-scaffold".

Production of the bTEBV

The scaffold material is pressed through the extruder (Figure 4). After passage of the material through the molder and submersing the extruded bTEBV in the hardening tube, filled with 2.5% [w/v] aqueous CaCl ₂ solution (Figure 4A), the bTEBV can be taken out and pulled over a stainless steel tube outlet (Figure 4B, D and G). The test solution (here stained in purple; Figure 4G) is then squeezed via the pressure syringe into the bTEBV. A close-up of the extruder with its central molder (Figure 4C), as well as of the readily fabricated bTEBVs, here with an inner diameter of 0.8 mm, is also shown (Figure 4E, F and G).

Through variation of the distance of the center of the central stab to the outer edge of the outlets of the laser-cut stainless steel aperture disc or through variation of the diameter of the central stab vessels with an outer diameter of 6 mm [inner diameter, 4 mm] (see Figure 7 right; for comparison, a commercial synthetic vessel graft is shown on the left) or 1.8 mm [inner diameter of 0.8 mm], as well as intermediate size vessels of 2.95 mm [0.95 mm] can be produced.

Analysis of the burst pressure of the bTEBV Human muscular arteries vary in size from an outer diameter of ~ 10 mm to about 0.5 mm (mean diameter ~ 4 mm and mean wall thickness ~ 1 mm), while the similarly sized veins have wider inner diameters, with a mean diameter of ~ 5 mm and a mean wall thickness of ~ 0.5 mm. Therefore the inventors have chosen for the burst pressure studies bTEBVs with an outer size of 6 mm [tube inner diameter of 4 mm] and 1.8 mm [central tube diameter of 0.8 mm]. The device used is shown in Figure 8A; for the experiments described here (37°C) the alternating flow rate was adjusted to 3-6 ml/min and the pressure was set between 25 and 100 mbar (18.7 and 75.0 mm Hg). Using this setting the vessels remained intact for over 14 days. During the pulse cycles the diameter of the vessels changed. As an example, the outer diameter of the smaller vessels expanded from 1.8 mm to 2.05 mm (Figure 8B to E).

In order to determine the burst pressure the primary pressure was increased, while the alternative flow rate remained unchanged. Under those conditions the burst pressure for the larger tubes was determined to be 850±155 mbar (n=20), while that for the smaller 1.8 mm tubes was 145±24 mbar. Those values are close to the ones measured for human saphenous veins which show a burst pressure at 2,200-3,000 mbar at a wall thickness of ~ 250 μιη. In a further comparison, the human arteries show a higher resistance and are characterized by a burst pressure of 2,700- 5,600 mbar at a wall thickness of 350-710 μιη.

Analysis of the resistance/hardness of the biomaterial

As a parameter for the elastic modulus the hardness of the N, O-CMC-polyP scaffold was determined. The newly-determined ferruled optical fiber-based nanoindenter was used as outlined under Methods. The elastic modulus of a biomaterial used for fabrication of the vessel graft should be close to the values measured for physiological cells. The elastic moduli of human arteries and veins are - 455 kPa. In comparison, graft types prepared from PET or ePTFE are considerably harder with a value of 1,900 kPa (Salacinski HJ, Goldner S, Giudiceandrea A, Hamilton G, Seifalian AM, Edwards A, Carson RJ. The mechanical behavior of vascular grafts: a review. J Biomater Appl 2001;15 :241-278) and 2,200 kPa (Zidi M, Cheref M. Mechanical analysis of a prototype of small diameter vascular prosthesis: numerical simulations. Comput Biol Med 2003;33:65-75), respectively. The elastic modulus of the vessels, prepared with the N, O-CMC-based biomaterial, described here, acquired a hardness of 475±61 (n=10) kPa after an incubation period of 5 min in 2.5% [w/v] CaCl ₂ solution. If 70% ethanol is added to the hardening solution consisting of 2.5% [w/v] aqueous CaCl ₂ the hardening process can be decisively accelerated (Figure 9). Even after a duration of stay of the bTEBV for 15 s the tubes reach an elastic modulus of 117±21 kPa (at 30°C), a value which increases during an incubation period of 180 s to 1,215±183 kPa. If the hardening process is performed at a lower incubation temperature than 30°C, at 10°C the hardening process proceeds slower, starting with only 84±9 kPa, but reaching after 180 s almost the same value as measured at 30°C. If the incubation temperature is increased to 50°C the kinetics of the hardening process is very much the same as measured for 30°C (Figure 9).

Effect of the adhesion-promoting oligo-/polymers on growth of HUVEC cells

The effect of the polycationic poly-L-lysine and poly-D-lysine as well as the His/Gly-tagged RGD on viability/growth of HUVEC cells was determined using the XTT assay. The cells, in 48 well plates, were incubated with the following scaffold samples: "control-scaffold", "poly(L-Lys)-scaffold", "poly(D-Lys)-scaffold" or "RGD-scaffold". After an incubation period of 7 d the amount of insoluble tetrazolium salt formed was determined. With respect to the "control-scaffold" all scaffolds supplemented with the adhesion-promoting compounds significantly increased the cell concentration from 0.87±0.11 ("control-scaffold"), to 1.79±0.19 ("poly(L-Lys)-scaffold"), 1.53±0.19 ("poly(D-Lys)-scaffold") and 2.89±0.24 ("RGD-scaffold"), reflecting that the His/Gly-tagged RGD peptide displays the strongest effect on cell growth (Figure 10).

The cell growth studies were paralleled with studies on the density of cells attached onto the surface of the scaffolds. The different types of scaffolds were incubated with HUVEC cells for 2 weeks under conditions described in Methods. After incubation the cells, attached to the surface of the scaffold, were fixed in formaldehyde vapor and stained for cytoskeleton structures, actin, with antibodies and for DNA in the nucleus with DRAQ5 (Figure 11). The data revealed that only a few clusters of cells are present on the surface of the "control- scaffold" (Figure 11 A, B; Figure IB), while a higher density is seen onto "poly(D-Lys)- scaffold" (Figure 11C, D; Figure 1C), and - even more - onto "poly(L-Lys)-scaffold" (Figure 9E, F), while the highest density is seen on scaffold containing the His/Gly-tagged RGD "RGD-scaffold" (Figure 11G, H; Figure ID). A co-addition of poly(L-Lys) to the His/Gly- tagged RGD-scaffold (poly(L-Lys)-scaffold/RGD-scaffold) is apparently not a better substrate for the HUVEC cells to attach to (Figure 1 II, J). Effect of the polycationic polymers on polyP caused increase of blood coagulation PolyP has been reported to enhance clot formation (Smith S A, Mutch NJ, Baskar D, Rohloff P, Docampo R, Morrissey JH. Polyphosphate modulates blood coagulation and fibrinolysis. Proc Natl Acad Sci USA 2006;103:903-908). This property of polyP could significantly affect the usefulness of the bTEBV as prosthetic vascular implants. Therefore the inventors determined the effect of the different matrices used for the fabrication of the bTEBV on the clotting time of human blood in vitro. The measurements were performed with whole blood contacted with equal amounts of matrices. After a contact period of 10 min the remaining erythrocytes in suspension were determined on the basis of their hemoglobin content. As summarized in Table 1, the presence of polyP in the basic scaffold [alginate-Ca ²⁺ -N, O-CMC- polyP] caused a significant reduction of the concentration of hemoglobin, as a measure for the free erythrocyte number, in the supernatant compared to the polyP-free control [alginate-Ca ²⁺ - NO-CMC]. Unexpectedly the inventors found that this effect is abolished in the presence of the polycationic polymers, poly(L-Lys), poly(D-Lys) and His/Gly-tagged RGD peptide, embedded into the alginate-Ca ²⁺ -N, O-CMC-polyP matrix. All polycationic polymers significantly inhibit the increase in on blood clotting caused by polyP if present in the alginate-Ca ²⁺ -N,0-CMC-polyP matrix [alginate-Ca ²⁺ -N, O-CMC-polyP matrix] (Table 1).

Table 1. Effect of the polycationic polymers, poly(L-Lys), poly(D-Lys) and His/Gly-tagged RGD peptide, embedded into the alginate-Ca ²⁺ -N, O-CMC-polyP matrix, on blood clotting rates, compared to controls without polyP [alginate-Ca ²⁺ -N, O-CMC] and with polyP [alginate-Ca ²⁺ -N,O-CMC-polyP]. The absorbance of hemoglobin from lysed noncoagulated erythrocytes was determined.

Application of the tube-like structures as neuronal sheets The neuronal sheets can be fabricated as follows. After or during the described preparation process of the inert hydrogel scaffold, consisting of the basic anionic polymers (modified chitosan and alginate) 2% of Na-polyP and 0.01% poly(L-Lys) are added. Using this material, tubes with an inner diameter of 0.55 mm have been prepared, as outlined above, applying the inventive extruder and the Ca ²⁺ hardening procedure. Then neuronal cells (rat cortical cells) were prepared using established techniques (Perovic S, Schleger C, Pergande G, Iskric S, Ushijima H, Rytik P, Miiller WEG. The triaminopyridine Flupirtine prevents cell death in rat cortical cells induced by N-methyl-D-aspartate and gpl20 of HIV-1. Europ J Pharmacol 1994;288:27-33). After purification of the neurons by treatment with cytosine arabinoside a >90% pure preparation, with respect to neurons, was obtained. Only ~ 10% GFAP -positive astrocytes were detected by using the established glial fibrillary acidic protein antibodies detection system. Those neurons were detached from the plastic culture dishes and filled into the tubes at a concentration of 2x1000 cells/ml; 50 μΐ of cell suspension were added to the tube opening. Then the 4 to 8 mm tubes were incubated under standard conditions (Perovic S, Schleger C, Pergande G, Iskric S, Ushijima H, Rytik P, Miiller WEG. The triaminopyridine Flupirtine prevents cell death in rat cortical cells induced by N-methyl-D-aspartate and gpl20 of HIV-1. Europ J Pharmacol 1994;288:27-33) for 72 h in a humidified 5%CO/95% air incubator at 37°C in minimum essential medium (MEM, Gibco), supplemented with heat- inactivated fetal bovine serum (Gibco). Then the tubes were opened by a lateral sectioning and the cultures were examined by phase-contrast light microscopy. A strong neurite outgrowth was observed with a length between 5 and 35 μΜ.

Application of the tube-like structures as light-waveguides

At first the light-waveguide was prepared by using the above described procedure and the matrix formed by the inert anionic polymers, Ν,Ο-CMC and alginate into which 10 μg/ml of Glu-tagged silicatein (Natalio F, Link T, Miiller WEG, Schroder HC, Cui FZ, Wang X, Wiens M. Bioengineering of the silica-polymerizing enzyme silicatein-a for a targeted application to hydroxyapatite. Acta Biomaterialia 2010;6:3720-3728) were added. This scaffold material was pressed through the inventive extruder. Diameter sizes of 0.15 to 0.20 mm were obtained. The fabricated gel-like tubes were injected into the CaCl ₂ aqueous solution. After an incubation period of 30 s the hardened tubes were collected with a forceps and incubated with 300 μΜ pre-hydrolyzed tetraethyl orthosilicate (TEOS) to obtain the ortho-silicate as the substrate for the silicatein. The resulting silica coated fibers obtained after silicatein reaction were adjusted/coupled in a holder that was placed immediately in front of the outlet of the laser light generator (Muller WEG, Link T, Schroder HC, Korzhev M, Neufurth M, Brandt D, Wang XH. Dissection of the structure-forming from the structure-guiding activity of silicatein: a biomimetic molecular approach to print optical fibers. J Mater Chem B 2014;2:5368-5377). A free-spaced green laser source (Hyperion Multi-Color Laser Source; XiO Photonics, Enschede; The Netherlands) at 532 nm with a power of 5-6 mW was used. The length of the fibers was between 0.5 and 3 cm. The amount of light input as well as the light penetrating through the fiber (at the output end of the fiber; the "optical light transmission") was measured with an optical spectrum analyzer (AQ-6315A from Ando Electronics, Kawasaki, Japan) as described (Muller WEG, Wendt K, Geppert C, Wiens M, Reiber A, Schroder HC. Novel photoreception system in sponges? Unique transmission properties of the stalk spicules from the hexactinellid Hyalonema sieboldi. Biosens Bioelectron 2006;21 : 1149-1155). The results revealed that the degree of light transmission decreased by ~ 5% during a penetration though 1 cm of light fiber and ~ 15 % during a path through 2 cm.

Methods

Materials

Sodium polyphosphate (Na-polyP of an average chain of 40 phosphate units) was used. The peptide NH ₂ -RGDGGRGDGGGRGDGGGHHHHHH-COOH (SEQ ID NO: 1) [named His/Gly-tagged RGD] had been chemically synthesized, applying the solid-phase peptide synthesis technique and using the Fmoc strategy.

The peptide NH 2 -DDDSQGEIQSDMAEEEDDDNVD-COOH (M r 2.47 kDa) (SEQ ID NO. 2), spanning the oligopeptide stretch aa_51 to aa_72 of silintaphin-2, was synthesized by solid phase peptide synthesis. The amino acid sequence of silintaphin-2 had been deposited at the EMBL/Genbank under the accession number FR681734.

Preparation of N, O-carboxymethyl chitosan

N, 0-carboxymethyl chitosan (N.O-CMC) can be prepared from chitosan (from shrimp shells) according to state-of-the-art procedures (Sakairi N, Suzuki S, Ueno K, Han SM, Nishi N, Tokura S. Biosynthesis of hetero-polysaccharides by Acetobacter xy/mwm-synthesis and characterization of metal-ion adsorptive properties of partially carboxmethylated cellulose. Carbohydrate Polymers 1998;37:409-414; Chen XG, Park HJ. Chemical characteristics of O- carboxymethyl chitosans related to the preparation conditions. Carbohydrate Polymers 2003;53:355-359; Chen SC, Wu YC, Mi FL, Lin YH, Yu LC, Sung HW. A novel pH- sensitive hydrogel composed of N, 0-carboxymethyl chitosan and alginate cross-linked by genipin for protein drug delivery. J Control Release 2004;96:285-300) and sterilized, for example, with a 12-W ultraviolet lamp (Syngene, Cambridge) at 254-nm wavelength (distance of 10 cm; 12 h). A suspension of 50 mg/ml of N,0-CMC, in sterilized physiological saline (0.9% [w/v] NaCl), supplemented with 30 μΜ silica (prepared from Na-silicate), is prepared and stirred until it becomes homogeneous. Then 20 mg/ml of solid Na-polyP is added and again stirred until a close to uniform solution is reached. The resulting hydrogel solution is supplemented with 50 mg/ml of sodium alginate and brought to homogeneity at 50°C, while stirring. If not mentioned otherwise this hydrogel, the "basic scaffold" is enriched with 0.1% low-melting gelatin (bovine) that is added together with the alginate.

Fourier transformed infrared (FTIR) spectroscopy is used in the attenuated total reflectance (ATR) mode to assure the substitutions of carboxymethyl groups at the amino group as well as the primary hydroxyl sites of the chitosan (FTIR-ATR; Varian 660-IR spectrometer with Golden Gate ATR auxiliary) (Chen SC, Wu YC, Mi FL, Lin YH, Yu LC, Sung HW. A novel pH-sensitive hydrogel composed of N, 0-carboxymethyl chitosan and alginate cross-linked by genipin for protein drug delivery. J Control Release 2004;96:285-300). Dried powder of sample is placed onto the ATR crystal directly. The spectra are acquired at 4000-750 cm ^"1 wave numbers with a 4 cm ^"1 resolution.

Preparation of O-carboxymethyl chitosan

O-carboxmethyl chitosan (O-CMC) can be prepared by reacting monochloroacetic acid with chitosan in isopropanol/NaOH solution using state-of-the-art procedures (e.g. Upadhyaya L, Singh J, Agarwal V, Tewari RP. Biomedical applications of carboxymethyl chitosans. Carbohydr Polym 2013;91 :452-466).

Preparation of N-carboxymethyl chitosan

N-carboxmethyl chitosan (N-CMC) can be obtained by reacting free amino groups of chitosan with glyoxylic acid and subsequent reduction of the resulting aldimine with sodium borohydride, as described (e.g. Upadhyaya L, Singh J, Agarwal V, Tewari RP. Biomedical applications of carboxymethyl chitosans. Carbohydr Polym 2013;91 :452-466).

Addition of polycations and His/Gly-tagged RGD It is crucial that a scaffold used to fabricate bTEBV provides a suitable template for the cells to adhere to. As shown in the Examples, the following components can be added to the basic scaffold at a concentration of 3 μg/ml; also higher and lower concentrations can be used poly(L-Lys) (for example as hydrobromide and with a mol wt 70,000-150,000), poly(D-Lys) (for example as hydrobromide and with a mol wt 70,000-150,000) or His/Gly-tagged RGD. Poly(L-Lys) and poly(D-Lys) can be added to the scaffold after hardening with CaCl ₂ , while His/Gly-tagged RGD can be added prior to the crosslinking of the basic scaffold components (Λ ,Ο-CMC, polyP, alginate and gelatin). Subsequently the material is extruded or printed and finally hardened as described below. The material is termed "poly(L-Lys)-scaffold", "poly(D- Lys)-scaffold" or "RGD-scaffold".

Fabrication of bTEBV: Extruder

The bTEBV are fabricated by an extruder through which the N,O-CMC/alginate/polyP/gelatin hydrogel is continuously pressed through a nozzle (Figure 3). The hydrogel is filled into a sterile 5 ml syringe (for example: Discardit II, BD Heidelberg). Prior to the extrusion process to fabricate the vessels, the syringe is centrifuged (in direction of the gravity) to remove the remaining air bubbles (3 min; 1500 rpm). Then the plunger is inserted and the filled syringe is stored at 4°C until use. At the beginning of the squeezing out step of the bTEBV material, the temperature of the syringe is brought to 20° C (30 min). The syringe is connected to the homemade extruder unit which contained a laser-cut stainless steel aperture disc. Three curved openings separated by 200 μιη wide partitions are arranged in this disc in a circular way in order to allow the passage of the hydrogel. The three hydrogel strands immediately fuse together after their release around a central stab. Variation of the distance of the center of the central stab to the outer edge of the outlets of the laser-cut stainless steel aperture disc as well as variation of the diameter of the central stab allows the fabrication of vessels with an outer diameter from 1.8 mm [comprising a central tube diameter of 0.8 mm] to 6 mm [tube inner diameter of 4 mm]; vessels with an intermediate size of 2.95 mm [0.95 mm] can also be produced. The tubes are immediately submersed into a 2.5% [w/v] aqueous CaCl ₂ solution. During a (routinely) 5 min incubation period the vessels formed acquire their final hardness. Before the experiments the ends of the vessels are clipped off; usually 6 cm long vessels are made.

In order to increase the hardness of the vessels the hardening procedure of the bTEBV is performed in a 2.5% [w/v] aqueous CaCl ₂ solution, supplemented with 70%> [v/v] ethanol. Under those conditions the duration of the hardening can be shortened to 3 min (in maximum).

Fabrication of tube-like structures serving as neuronal sheets

The inventive method can be used for fabrication of neuronal sheets as follows. The inert hydrogel scaffold consists of the anionic polymers N,O-CMC and alginate. The polyanion Na- polyP (final concentration, 2%) and the polycation poly(L-Lys) (final concentration, 0.01%) are added either after or during the preparation process, as described above. The inventive extruder and the aforementioned Ca ²⁺ hardening procedure are used. According to this procedure, tubes with various diameters can be generated; in the experiment described in the Examples, tubes with an inner diameter of 0.55 mm have been used.

Fabrication of tube-like structures serving as light-transmitting fibers

The silica-coated light-transmitting fibers are prepared by using the aforementioned inventive method. The inventive extruder is used through which the inert anionic polymers, N, O-CMC and alginate, supplemented with 10 μg/ml of recombinant Glu-tagged silicatein are pressed. Using this method, gel- like tubes with various diameter sizes, for example 0.15 to 0.20 mm, can be obtained, which are injected into the CaCl ₂ solution; the incubation period is approximately 30 s. The hardened tubes are then incubated with a suitable silicatein substrate, e.g., 300 μΜ pre-hydrolyzed tetraethyl orthosilicate (TEOS).

The recombinant Glu-tagged silicatein can be prepared, for example, from the short form of the mature silicatein-a of Suberites domuncula as described (Natalio F, Link T, Miiller WEG, Schroder HC, Cui FZ, Wang X, Wiens M. Bioengineering of the silica-polymerizing enzyme silicatein-a for a targeted application to hydroxyapatite. Acta Biomaterialia 2010;6:3720- 3728). The protein can be purified, for example, by Ni ²⁺ -NTA affinity chromatography and unfolded, for example, in 6 M urea/5 mM imidazole and finally refolded in 50 mM Tris/HCl buffer (0.5 M L-arginine, glutathione [9 mM glutathione / 1 mM oxidized glutathione] redox couple, 0.3 M NaCl, 1 mM KC1).

Bioprinting of gel discs

Discs with the material used for the testing of adhesion and growth of endothelial cells are bioprinted, following described procedures (Neufurth M, Wang XH, Schroder HC, Feng QL, Diehl-Seifert B, Ziebart T, Steffen R, Wang SF and Miiller WEG. Engineering a morphogenetically active hydrogel for bioprinting of bioartificial tissue derived from human osteoblast-like SaOS-2 cells. Biomaterials 2014; 35:8810-8819); for example, a 3D- Bioplotter, 4th generation blotter, from Envisiontec can be used. The respective hydrogel preparation is filled into sterile 30 ml printing cartridges (for example: Nordson EFD) and centrifuged for 3 min at 1500 rpm to remove air bubbles. After connecting a 0.25 mm tapered polyethylene printing tip (for example: Nordson EFD, Pforzheim; Germany) the cartridge is placed into the preheated (25°C) printing head of the bioplotter. At 25°C, using a pressure of 1.5 bar and a printing speed of 16 mm/s cylindrical scaffolds measuring 7.5 x 0.4 mm are printed as described (Neufurth M, Wang XH, Schroder HC, Feng QL, Diehl-Seifert B, Ziebart T, Steffen R, Wang SF and Miiller WEG. Engineering a morphogenetically active hydrogel for bioprinting of bioartificial tissue derived from human osteoblast-like SaOS-2 cells. Biomaterials 2014; 35:8810-8819). The strand distance between the printed cylinders is set to 1 mm resulting in a pore size of the printed layers of approximately 0.5 x 0.5 mm. Those scaffolds are printed directly into sterile 94 mm Petri dishes, supplemented with 2.5% [w/v] CaCl ₂ as crosslinking solution, as indicated with the results. The printed discs with a diameter 9.5 mm can be hardened in 2.5% [w/v] CaCl ₂ for crosslinking for 5 min. During this treatment the discs shrink to 7.5 mm.

Burst pressure experiments

The studies to determine the burst pressure of the fabricated bTEBV can be performed using tubes of a size of 1.8 mm [inner diameter: 0.8 mm] and 6 mm [4 mm]. Those tubes are pulled on both ends over either FEP (fluorinated ethylene propylene) hoses from intravenous catheters, measuring 0.9 mm outer diameter or 4 mm luer-hose-connectors (Carl Roth GmbH). The hoses are inserted approx. 10-20 mm deep into the lumen of the bTEBVs. This setup is connected to a pump system (for example: BioRad Econo Pump) with an attached pressure gauge. The liquid reservoir is placed into a water bath at 37°C to mimic body temperature. For routine experiments the pump is adjusted to 3-6 ml/min and the pressure is regulated between 25-100 mbar using a hose clamp, mounted downstream of the sample. For the determination of the burst pressure, the liquid flow is totally blocked downstream of the sample. The pressure is then steadily increased until the sample shows severe leakage or complete burst.

Elastic modulus, reduced Young's modulus, of the biomaterial As a parameter for the elastic modulus the hardness of the N, O-CMC-polyP scaffold can be determined, for example, by using a ferruled optical fiber-based nanoindenter as described (Chavan D, Andres D, Iannuzzi D. Note: ferrule-top atomic force microscope. II. Imaging in tapping mode and at low temperature. Rev Sci Instrum 2011;82:046107; doi: 10.1063/1.3579496; Chavan D, van de Watering TC, Gruca G, Rector JH, Heeck K, Slaman M, Iannuzzi D. Ferrule-top nanoindenter: an optomechanical fiber sensor for nanoindentation. Rev Sci Instrum 2012;83:115110; doi: 10.1063/1.4766959). The indents are depth controlled (10 μιη) and the loading and unloading period is set to 2 s. Based on the load-displacement curves the reduced Young's modulus [RedYM] is calculated.

Effect of blood clotting time

The influence of the polycationic polymers, poly(L-Lys), poly(D-Lys) and His/Gly-tagged RGD peptide, embedded into the alginate-Ca ²⁺ -N, O-CMC-polyP matrix, compared to the control without polyP [alginate-Ca ²⁺ -N, O-CMC] and with polyP [alginate-Ca ²⁺ -N, O-CMC- polyP], on blood clotting time can be determined, for example, by the assay described by (Shih MF, Shau MD, Chang MY, Chiou SK, Chang JK, Cherng JY. Platelet adsorption and hemolytic properties of liquid crystal/composite polymers. Int J Pharm 2006;327: 117-125). Equal amounts of samples (110 mg) are submersed in bottles placed in a thermostated water bath at 37°C for 10 min. Then 300 μΐ of human blood sample (acid-citrate-dextrose with 20 μΐ/ml of 100 mM CaCl ₂ ) is dropped on the surface of the matrices until they are completely covered. Then the assays are continued to be incubated (37°C) for 10 min. Then 15 ml of distilled water are added without disturbing the clotted blood. Subsequently 10 ml aliquots are taken, centrifuged (100xg; 30 s) and the supernatant is collected and the clotting test is performed spectrophotometrically at 542 nm.

Endothelial cell culture

As a model for endothelial cells, for example, human umbilical vein/vascular endothelial cells, the transformed HUVEC cells, EA.hy 926 ATCC CRL-2922 can be used. The cells are grown in Dulbecco's Modified Eagle's Medium (DMEM; low glucose), supplemented with 10% fetal bovine serum (FBS), 50 μg/ml gentamycin, 2 mM glutamine and 2% HAT Media Supplement. The cells are layered onto the printed gel discs at a density of 10,000 cells/cm ² in 48 well plates (0.95 cm ² growth area; Greiner). Splitting of the medium/serum is performed twice in a week. Growth of the cells, after an incubation period of 7 d, can be determined, for example, by a colorimetric method based on the tetrazolium salt XTT. The printed discs with a diameter 1.1 cm are hardened in 2.5% [w/v] CaCl ₂ for crosslinking for 5 min. The following discs have been used for the experiments: "control-scaffold", containing none of the adhesion promoting polymers poly(L-Lys), poly(D-Lys) or RGD-scaffold, "poly(L-Lys)-scaffold", "poly(D-Lys)- scaffold" or "RGD-scaffold". A 300 μΐ cell suspension (80,000 cells/ml) is added per well and incubation is continued for 7 d, followed by determination of viability using the XTT assay.

Staining of endothelial cells

The cells attached to the surface of the scaffold can be fixed, for example, in formaldehyde vapor, then blocked (15% blocking solution and 0.1% Triton X-100 in PBS) for 1 h (room temperature). To visualize the cells, they can be stained with rhodamine phalloidin (for example: PHDR1; Biomol) as well as, for example, DRAQ5 (Biostatus Ltd; nuclear stain). The actin complexes are visualized at an excitation of 550 nm and an emission of 570 nm, while for detection of the DRAQ5 fluorescence the cells are excited at 635 nm, and recorded at an emission of 705 nm.

Primary neuronal cell culture

Rat cortical cell cultures can be prepared, for example, from the brains of Wistar rat embryos (18 days old) as follows. After isolation cerebral hemispheres are placed into growth medium (Dulbecco's modified Eagle's medium, supplemented with 30 mM glucose, 24.5 mM KC1, 7 μΜ /?-aminobenzoic acid, 100 mU/1 of insulin, 14 mM NaHC0 ₃ and 5 mM Hepes). Brain tissue is dissociated in medium using 0.025%) of trypsin (10 min; 37°C). The proteolytic reaction is stopped by addition of trypsin inhibitor. After sedimentation, the single cell suspension is centrifuged. The pellet containing dissociated neuronal cells is resuspended in growth medium, containing 10% fetal calf serum and placed into 25 cm ² plastic flasks (Nunc) (2.5 x 10 ⁵ cells/cm ² ) that had been coated with poly(L-Lys) (5 μg/ml; 3 ml). The cells are kept in an atmosphere of 95% air and 5% C0 ₂ . To enrich the percentage of neurons the culture is incubated for 48 h in the presence of 10 μΜ cytosine arabinoside. After incubation for 24 h in the cytosine arabinoside- free medium/serum the cells are used for the experiments. The cultures contain > 90% of neurons and ~ 10% glial fibrillary acidic protein (GFAP)-positive astrocytes and macrophages.

Light microscopic analyses Digital light microscopic studies can be performed, for example, using a VHX-600 Digital Microscope (Keyence) equipped with a VH-Z25 zoom lens. The fluorescence of the cells is excited at 635 nm, and the emission is recorded at 705 nm in an Olympus 1X71 fluorescence microscope.

Exposure of the silica-coated fibers to laser light

The light-transmitting silica-coated fibers fabricated by the inventive method are inspected under a light microscope (for example, an Olympus AHBT3 microscope) with an attached CCD-Camera. The fibers are coupled with a free-spaced green laser source, for example, a Hyperion Multi-Co lor Laser Source (XiO Photonics) at 532 nm with a power of 5-6 mW. The transmitted light can be measured, for example, with an optical spectrum analyzer (e.g., AQ- 6315 A; Ando Electronics, Kawasaki; Japan) as described (Miiller WEG, Wendt K, Geppert C, Wiens M, Reiber A, Schroder HC. Novel photoreception system in sponges? Unique transmission properties of the stalk spicules from the hexactinellid Hyalonema sieboldi. Biosens Bioelectron 2006;21 : 1149-1155).

Statistical analysis

The results are statistically evaluated using paired Student's t-test.