EP3638327A1

EP3638327A1 - Method for producing multilayered composites and their uses

Info

Publication number: EP3638327A1
Application number: EP18734974.1A
Authority: EP
Inventors: Juan Pablo ACEVEDO COX; Maroun Khoury; Tamara Lissette AKENTJEW FAURE
Original assignee: Cells For Cells SA
Current assignee: Cells For Cells SA
Priority date: 2017-06-16
Filing date: 2018-06-18
Publication date: 2020-04-22
Also published as: WO2018229734A1

Abstract

The present invention provides a method for producing a multilayered composite, the multilayered composite produced by the method as well as the use of the multilayered composite to replace or patch a number of tissues. The methods of the present invention provide a multilayered composite that can have similar mechanical properties to that of native tissue.

Description

DESCRIPTION

METHOD FOR PRODUCING MULTILAYERED COMPOSITES AND THEIR USES

Technical field

The present invention can be included in the field of medicine and tissue engineering. Specifically, the present application relates to a multilayered composite, a method of producing the composite and the uses of the composite for patching and/or replacing tissues.

Background art

A large number of strategies have been proposed in order to solve the issue of vascular implant availability. For many years synthetic polymers were used to fabricate grafts that have shown great success in replacing large caliber vessels. Nevertheless, these implants have presented poor results for diameters smaller than 6 mm, presenting intimal hyperplasia and thrombosis. Nowadays there is a consensus among the scientific community that all designs should contain an endothelial layer or elements that mimic endothelial layer functionality in order to avoid post- surgery issues and extend their permeability [1 , 2]. Some of the new advances in small diameter vascular grafts are listed below, categorizing them according to fabrication strategies:

Synthetic grafts: These are manufactured using polymers like Dacron (Polyethylene terephthalate), Teflon (Polytetrafluoroethylene) and Gore-Tex (stretched polytetrafluoroethylene or ePTFE). They do not possess bioactivity and the cell adhesion is very limited. Additionally, material degradation might create acidic microenvironments in the surrounding area, affecting cell functionality [3]. These grafts are commonly highly susceptible to infections. To improve their performance, endothelial cell attachments have been stimulated by internal coating of agents such as cell-adhesion peptides [4] or sulfated silk fibroin [5]; alternatively, the grafts have also been coated with anticoagulant components such as heparin. Nevertheless, these products continue exhibiting issues associated with partial endothelization, such as hyperplasia and thrombosis, and mechanical incompatibility and fatigue. A research done by Klinkert and collaborators showed that even for large-diameter vessels, only 39% of 5 year old synthetic grafts maintained an adequate permeability [6]. Lastly, another drawback is that they cannot be remodeled by the surrounding or growing tissues, making them inadequate for young patients. Grafts manufactured with synthetic biodegradable scaffolds: Yue and collaborators were able to fabricate porous scaffolds from degradable polylactic acid (PLA), over which they seeded smooth muscle cells and applied them as vascular grafts in rats. Results indicated that cellularized scaffolds had a better performance than decellularized scaffolds [7]. Other researchers used other materials such as polyglycolic acid (PGA) [8] and PGA with polyhydroxyialkanoate [9] and obtained similar results. However, it must be mentioned that none of these materials applied as vascular grafts had mechanical properties approximating the natural vessels.

Grafts manufactured with natural scaffolds: Different natural scaffolds can be used as grafts, such as autograft from saphenous vein, cryopreserved umbilical cords or decellularized tissue.

Allogenic Transplants: The first reports of this technique indicated that these grafts presented ruptures shortly after surgery [10, 11 ], which limited their use. Nevertheless, stabilizing technologies have considerably improved their performance [12], and this option is now used for second vascular bypass surgeries in severe cases. Decellularized scaffolds: Researchers have successfully explored the use of decellularized tissues as scaffolds for manufacturing vascular grafts. Using these scaffolds, the resulting constructs have adequate structural strength and mechanical properties granted by the natural extracellular matrix (ECM) of the tissue. Among the natural scaffolds that have been successfully studied are rat arteries [13], pig abdominal aorta [14], human amniotic membrane [15], umbilical arteries and canine carotid arteries [16]. However, this approach has the disadvantage of restricted dimensional and mechanical options, and importantly cell seeding remains a limiting step [17]. Additionally, there is a latent risk of pathogen transmission to the patient.

The work done by Prof. Niklason takes a novel approach on the use of decellularization and recellularization of fabricated scaffolds [18]. Her manufacturing procedure uses human or canine cells seeded on a PGA cylindrical scaffold. During culture, cells deposit new ECM while the PGA is degraded over time. Afterwards the resulting tissue is decellularized to obtain a new non- immunogenic scaffold composed of ECM proteins. This descellularized scaffold can be used directly for bypasses larger than 6 mm in diameter. Experiments with baboons have shown good results in a restricted cohort. For grafts smaller than 6 mm in diameter, it has been established that recellularization with patients' endothelial cells (EC) is required. However, only low level of EC internal coverage is achieved (14%) and coronary bypass experiments have not been followed up for longer than a month [15]. One important practical disadvantage is that this type of graft requires about 5-10 month of manufacturing, restricting importantly the mass use and commercialization. Vessels constructed without a scaffold: ECM components such as collagen, laminin, fibronectin, among others can be mixed with cells before polymerization or crosslinking. Hirai and collaborators manufactured a blood vessel by pouring a blend of canine smooth muscle with collagen in a cylindrical cast where the solution was polymerized [19]. This structure was then endothelized and reinforced with a Dacron outer mesh. This technique allows fabrication of vessels of different sizes depending on the selected mold, and mechanical characteristics by altering the amount of cells, collagen and incubating time.

Another technique used for generating cylindrical structures is the dipping technology where a rod is immersed in a pre-polymerized solution and cross-linked during or after the rod is retrieved. This option has been explored by Prof. Kaplan who constructed a one layer cylindrical structures with silk fibers that was subsequently seeded with cells [20]; on the other hand, Prof. Khademhosseini generated a thin cell-laden layer with a cylindrical shape by dipping a metal rod into a prepolymerized solution of Gelatin and cells [21], however, he did not explore the feasibility of this particular technique in generating more complex multilayer vessel-like structures. Our Tissue Engineering group at Universidad de los Andes has advanced the dipping technology, and explored its applicability in fabricating complex multilayer cylindrical structures, combining cellularized layers and elastomeric tough biomaterials. We have developed a computer numerically controlled (CNC, WO2017/064667) machine that automatizes the dipping process, adding a simultaneous rotational movement to organize the microstructural alignment of deposited biomaterials. To date, the research group has encountered results regarding the standardization of dimensional features of the multilayer cylindrical structures, the use of different biomaterials, and the precise inclusion of cellularized layers within the vessellike construct. Finally, Dr. L'Hereux created a method that uses a cell sheet of fibroblasts [22], smooth muscle cells or combination of both [23], which are subsequently detached and rolled using a metal rod. After long and expensive manufacturing process, including cell seeding, rolling and maturation, this method creates a cylindrical structure capable to withstand blood pressure. Disadvantage related to this approach are: difficult automatization and cost-effective production, and standardization of mechanical properties and safety are difficult due to the cell donor heterogeneity.

Supporting meshes: Electrospinning technology allows the deposition of thin intricate mesh of nanofibers over a cylindrical structure, which can mimic, to a certain extent, the vessels' ECM. This electrospun fiber-based layers allows cell seeding, adhesion and correct cell orientation following the fibers alignment. Among the biopolymer successfully used with this technique, we can find poly-e-caprolactone (PCL) [24], chitosan/PCL [25], Poly(ester urethane)urea [26], poly (diol citrate)-collagen-proteoglycan [3], amongst many others. One of the advantages of this technology is that mechanical properties of natural blood vessels can be reproduced using electrospinning fabrication. However, it requires the generation of strong electrical field in the process, limiting the possibility of combination with other fabrication technique or cells. Our associated principal investigator from Manchester University, Prof. Jonny Blaker, has developed an alternative technology named Solution Blow Spinning (SBS) to generate similar structures [27]. SBS is a rapid sub-micron/nanofibre production technique that can produce nanofibres circa 100 times faster than electrospinning and does not require any complicated equipment or electric fields. Simply a polymer solution (or thermoplastic polymer at elevated temperature in the case of melt spinning) is injected. We envisioned that this technique can be combined with our dipping-rotation process to developed vessel-like structures with cellularized multilayer and mechanical properties similar to natural vasculature, and capable to resist suturing and rupture by blood pressure.

Mesenchymal Stem Cells and Vessel graft Engineering: Mesenchymal Stem Cells (MSCs) are capable of self-renewal and can differentiate to cell types such as osteocytes, chondrocytes, smooth muscle and endothelium. Some of their advantages are related to high availability and proliferation rate, harvesting through minimally invasive procedures from tissues such as bone marrow, adipose tissue, dental pulp, etc., and more importantly, they present a very low immunogenicity, possess immunosuppressing properties and anti-thrombogenic activity [28-30].

MSCs have been used in new vessel graft designs. Mirza and collaborators seeded MSCs over a synthetic graft, and observed partial differentiation to smooth muscle and endothelium after transplantation which improved the graft's patency [31]. Cho and his research team used differentiated MSCs to seed decellularized canine scaffolds, showing improved patency and potential to be used in small-diameter vessel fabrication [16]. Lastly, in Switzerland a 10 year old patient was successfully transplanted with an allogenic decellularized and recellularized vessel using her own MSCs extracted from her bone marrow[32].

There is currently a need for composites which have similar properties to native tissue. Therefore, it is an objective of the present invention to provide composites which have similar mechanical properties and which could be used to patch of replace native tissues.

Figures Figure 1: Strain-stress curves of polycaprolactone (PCL) sub-layers in longitudinal direction of 8 cycles of fibers deposited at 67° (black) and fibers deposited at 21 ° (grey) without preconditioning. Figure 2: Strain-stress curves of a) Methacryloyl gelatin-alginate (GEAL) layers reinforced with fibers at 21 °, b) preconditioned GEAL layers reinforced with fibers at 21°, c) preconditioned GEAL layers reinforced with wavy fibers at 21 °.

Figure 3: Microscopic images at 10X of a) fibers deposited at 21° and b) fiber deposited with waviness at 21 °.

Figure 4: Stress-strain curves of vascular graft layers fabricated from GEAL (also referred to as "GelBMa") sublayers, reinforced with PCL sublayers. Testing of different compositions allowed us to approach the curve of native media (grey dotted line) and adventitia (grey line) of coronary arteries (Holzapfel, 2005). Fibers deposited at 67° using a manufacture processing composed of 10 cycles of PCL sublayer intercalated with 9 dipping/crosslinking of GEAL solution (solid circle), 8 cycles of PCL sublayer intercalated with 7 dipping/crosslinking in GEAL solution (solid square) and 5 cycles of PCL sublayer intercalated with 15 dipping/crosslinking in GEAL solution (solid diamond). Fibers deposited at 21° with a total composition of 10 cycles of PCL sublayer intercalated with 4 dipping/crosslinking in GeAL solution (open circle), 8 cycles of PCL sublayer intercalated with 6 dipping/crosslinking in GEAL solution (o square) and 4 cycles of PCL sublayer intercalated with 8 dipping/crosslinking in GEAL solution (open diamond).

Figure 5: Stress-strain curves of outer and middle layer of GEAL reinforced with PCL fibers vascular graft. For outer layers the fibers were deposited at 67° (line) and human coronary artery media layer model response (grey dotted line) with the range (light grey) in a) longitudinal and b) circumferential directions. For middle layers the fibers were deposited at 21 ° (line) and human coronary artery media layer model response (grey dotted line) with the range (light grey) in c) longitudinal and d) circumferential directions.

Figure 6: Stress-strain curves of GEAL reinforced with PCL fibers vascular (line) and human coronary artery media layer model response (grey dotted line) with the range (light grey) in a) longitudinal and b) circumferential directions, (c) Cyclic tensile testing in circumferential direction.

Figure 7: D/DO vs pressure curves for GEAL reinforced with PCL fibers vascular graft (line) compared with human coronary arteries (solid circle) (Claes, 2010; van Andel, 2003) at three different values of axial prestretch. a) e_z=10% of axial prestretch.b) e_z=20% of axial prestretch. c) e_z =25 % of axial prestretch.

Figure 8: Cell density and distribution in GEAL reinforced with PCL fibers vascular grafts, a) Fluorescence microscopy image of a transversal cut of a vascular graft on day of manufacturing, b) Cell density at 1 and 7 days after cell encapsulation. Figure 9 shows the comparison between native skin and our skin graft. A J-shaped response similar to the native skin is observed in the skin graft stress stain curve for both 3 and 15 cycles. Summary of the invention

The present invention provides a method for producing a multilayered composite comprising the steps of (a) dipping a template into a pre-polymerized solution comprising gelatin and a photo- initiator, wherein the gelatin is chemically functionalized to become reactive to polymerization or cross-linking in the presence of free radicals, (b) exposing the pre-polymerized solution attached to the template to a wavelength of light which stimulates the photo-initiator and causes the gelatin to polymerize or cross-link, (c) optionally repeating steps (a) and (b) to obtain the desired number of layers, (d) depositing a fiber layer on the template at an equal or opposite angle to a naturally occurring fiber angle, wherein the template is rotated or moved so that the fibers are deposited in a wavy pattern, (e) optionally repeating step (d) to obtain the desired number of layers, (f) optionally repeating steps (a) to (e) to obtain the desired number of layers, and (g) preconditioning the resultant multilayered composite by stretching and relaxing it. Further, the present invention provides multilayered composite obtained or obtainable through any one of the methods provided and the use of the multilayered composite for the replacement or patching of blood vessels, skin, cartilage, tendons, ligaments, fistulas, stomach, esophagus, intestines, uterine tubes, larynx, urethra or nerve guidance conduits.

Detailed description of the invention

Definitions

The terms "individual", "patient" or "subject" are used interchangeably in the present application and are not meant to be limiting in any way. The "individual", "patient" or "subject" can be of any age, sex and physical condition.

The term "multilayered composite" refers to any composition which comprises at least two layers wherein at least one layer is a hydrogel layer and at least one layer is a fiber layer. Preferably the multilayered composite is biocompatible and it may comprise synthetic and/or natural components. An example of a natural component is gelatin which is found in nature. An example of a synthetic component is polycaprolactone which is man-made. In a preferred embodiment, the multilayered composite is a biomaterial.

The term "biocompatible" refers to any material which is not harmful or toxic to living tissue.

The term "hydrogeT refers to a network of polymer chains that are hydrophilic, sometimes found as a colloidal gel, in which water is the dispersion medium.

The term "fiber" refers to a natural or synthetic substance that is significantly longer than it is wide.

The term "wavy" refers to any pattern which is characterized by consisting of a series of undulating and wave-like curves.

The term "mesh" refers to any structure in which fibers loop or knot at intersections resulting in a structure with open spaces between the fibers. For an example of a mesh, see Figure 3B of the present application.

The term "naturally occurring fiber angle" refers to an angle with respect to a defined axis at which the arrangement of natural fibers aligns in natural tissues, such as blood vessels (in this case, the angle is usually respect to the circumferential axis). Alternatively, when physical and experimental evidence is not available to define the reference angles, "naturally occurring fiber angle" can be determined for a soft tissue using Holzapfel's model [33]. Holzapfel's model represents the strain-energy function of a tissue layer, which requires parameters related to the material components and geometrical aspects. Concerning material parameters, the model includes the isotropic deformation component, usually not related to the fiber component, and the anisotropic deformation component, provided basically by the reinforcing fiber component. Geometrical parameters considered in the model are the layer thickness and the fiber angle with respect to the defined axis. In order to identify the material parameters, they are fitted to the experimental Cauchy stress data of a tissue layer using the non-linear Levenberg-Marquadt algorithm. With these fitting processes and the model, one can calculate the angle at which the collagen fibers are oriented in a tissue. In 2005, Holzapfel applied this model to determine the naturally occurring fiber angle of the two layers of a coronary artery, media and adventitia layer [34].

The term "gelatin" refers to a hydrolyzed form of collagen, wherein the hydrolysis results in the reduction of the protein fibrils into its constituent polymer chains.

The term "photo-initiator" refers to a compound that undergoes a photoreaction on absorption of light, producing reactive species. Examples include: benzoyl peroxide, azobisisobutyronitrile, camphorquinone, irgacure and darocure. The term "template" refers to a structure composed of any material which does not interfere with the formation of the multilayered composite and which does not leave toxic or harmful substances in the multilayered composite. The structures may be of any shape or size and its features will depend on whatever size or shape the resultant multilayered composite should have. In a preferred embodiment, the template is a flat surface or a rod. Preferably, the template is a rod. Method

In a first aspect, the present invention provides a method for producing a multilayered composite comprising the steps of (a) dipping a template into a pre-polymerized solution comprising gelatin and a photo-initiator, wherein the gelatin is chemically functionalized to become reactive to polymerization or cross-linking in the presence of free radicals, (b) exposing the pre-polymerized solution attached to the template to a wavelength of light which stimulates the photo-initiator and causes the gelatin to polymerize or cross-link, (c) optionally repeating steps (a) and (b) to obtain the desired number of layers, (d) depositing a fiber layer on the template at an equal or opposite angle to a naturally occurring fiber angle, wherein the template is rotated or moved so that the fibers are deposited in a wavy pattern, (e) optionally repeating step (d) to obtain the desired number of layers, (f) optionally repeating steps (a) to (e) to obtain the desired number of layers, and (g) preconditioning the resultant multilayered composite by stretching and relaxing it. Steps (a) to (c) may occur before or after steps (d) and (e). Steps (a) and (b) may be repeated several times depending on how many hydrogel layers (i.e. layer which comprises gelatin) the skilled person wishes to have. Preferably, steps (a) and (b) are repeated 1 to 100 times. More preferably, 1 to 20 times. Step (d) may be repeated several times depending on how many fiber layers the skilled person wishes to have. Preferably, step (d) is repeated 1 to 100 times. More preferably, 1 to 100 times. Steps (a) to (c) and steps (d) and (e) may occur in an intercalated fashion so that the resultant composite has alternating layers of hydrogel and fibers.

In a preferred embodiment, the pre-polymerized solution comprises a UV photo-initiator which leads to cross-linking of the functionalized gelatin after exposure to UV light. The pre-polymerized solution is a solution comprising a polymer which has not yet been polymerized or crosslinked. More specifically, the pre-polymerized solution comprising functionalized gelatin is a solution comprising functionalized gelatin. The amount of functionalized gelatin can be in the range of 1 -20% w/v, preferably 8- 12% w/v and even more preferably 10% w/v based on the total amount of pre-polymerized solution.

The functionalized gelatin is a gelatin, the amino acidic chain of which is functionalized using a chemical agent which comprises a chemical group consisting of methacryloyl groups, acryloyl groups or any functional group or a moiety capable of mediating formation of a polymer or reaction with a surface or other molecule. Functional groups include the various radicals and chemical entities taught herein, and include alkenyl moieties such as acrylates, methacrylates, dimethacrylates, oligoacrylates, oligomethacrylates, ethacrylates, itaconates or acrylamides. Further functional groups include aldehydes.

Other functional groups may include ethylenically unsaturated monomers including, for example, alkyl esters of acrylic or methacrylic acid such as methyl methacrylate, ethyl methacrylate, butyl methacrylate, ethyl acrylate, butyl acrylate, hexyl acrylate, n-octyl acrylate, lauryl methacrylate, 2-ethylhexyl methacrylate, nonyl acrylate, benzyl methacrylate, the hydroxyalkyl esters of the same acids such as 2-hydroxyethyl acrylate, 2-hydroxyethyl methacrylate, and 2-hydroxypropyl methacrylate, the nitrile and amides of the same acids such as acrylonitrile, methacrylonitrile, and methacrylamide, vinyl acetate, vinyl propionate, vinylidene chloride, vinyl chloride, and vinyl aromatic compounds such as styrene, t-butyl styrene and vinyl toluene, dialkyl maleates, dialkylitaconates, dialkyl methylene-malonates, isoprene, and butadiene. Suitable ethylenically unsaturated monomers containing carboxylic acid groups include acrylic monomers such as acrylic acid, methacrylic acid, ethacrylic acid, itaconic acid, maleic acid, fumaric acid, monoalkylitaconate including monomethyl itaconate, monoethylitaconate, and monobutylitaconate, monoalkyl maleate including monomethyl maleate, monoethyl maleate, and monobutyl maleate, citraconic acid, and styrene carboxylic acid. Suitable polyethylenically unsaturated monomers include butadiene, isoprene, allylmethacrylate, diacrylates of alkyl diols such as butanedioldiacrylate and hexanedioldiacrylate, divinyl benzene, and the like. It is preferred that the amino acidic chain is functionalized with methacryloyl groups to give methacrylated gelatin. More preferably, the amino acidic chain is functionalized by using a chemical agent which provides methacryloyl, methacrylamide, acrilamide and/or acryloyl functionalization at the amino acid side chain of polymers.

The degree of functionalization of the acidic side chain of the gelatin polymer with a chemical agent capable of polymerizing or crosslinking in presence of free radicals is from 10% to 100%, preferably from 20% to 100%, more preferably from 30% to 100%, more preferably from 40% to 100%, more preferably from 50% to 100%, more preferably from 60% to 100%, more preferably from 70% to 100%, more preferably from 80% to 100%, more preferably from 90% to 100%. The amino acids involved in the functionalization can be one or more selected from the group consisting of serine, threonine, arginine tyrosine, lysine and others. Preferably, the functionalized amino acid is the lysine residue.

Usually, the temperature of the pre -polymerized gelatin solution in step (a) is from 26 to 40 °C, preferably from 28 to 37 °C and more preferably 37 °C. Different temperatures can also be used depending on the type of the functionalized gelatin used. Indeed, as specified throughout the description the pre-polymerized composition comprising gelatin has to be a solution. Thus, the skilled person will be able to set the temperature without undue burden.

In a preferred embodiment, wherein the gelatin of step (a) is modified using a chemical agent which provides methacryloyl and/or acryloyl functionalization at the amino acid side chain of polymers. Preferably, the chemical agent is methacrylic anhydride.

Suitable amounts of photo-initiator in the pre-polymerized solution are from 0.1 to 10% w/v. In some cases the amount of photo-initiator is from 0.1 to 1 % w/v, more preferably from 0.2 to 0.5% w/v.

In a preferred embodiment, the pre-polymerized solution can further comprise alginate or salts thereof or derivatives thereof. Suitable salts include, for example, sodium, potassium and lithium. The presence of these compounds can induce variation in the viscosity of gelatin which in turns influences the thickness of the layers. Hence, by selecting the amount of alginate in the pre-polymerized solution it is possible to increase and select the desired thickness of the layers.

Variations in the viscosity of gelatin solution can also be achieved by: (i) changing the degree of collagen hydrolysis during the process of gelatin extraction, (ii) using a different concentration of gelatin in solution, (iii) maintaining the gelatin solution at a different temperature, and varying the degree of partial polymerization or cross-linking of constituent polymer chains. Through any one of these approaches, the thickness of the layer can be further controlled. The pre-polymerized solution can comprise from 0.005% to 5% w/v alginate or salts thereof or derivatives thereof, preferably from 0.15% to 2% w/v, even more preferably from 0.15% to 1.5% w/v or from 0.3% to 0.6% w/v. The pre-polymerized solution according to the present invention can further comprise at least one compound selected from the group consisting of gelatin, chitosan, gellam gum, collagen, elastin, cellulose mixtures thereof, salts thereof and derivatives thereof.

In a preferred embodiment, the pre-polymerized solution comprises viable cells, proteins, extracellular vesicles, genetic material or polynucleotides, drugs and/or polymeric particles. Polymeric particles are micro or nano particles that can contain other elements for control release of those. Basically, a particle is a nano or micro-bead or particles compose of any polymeric compound capable to encapsulate other type of compounds, typically with biological activity. The viable cells may be stem cells or fully differentiated cells. Preferably, the viable cells are chosen from a list consisting of mesenchymal stem cells, endothelial cells, smooth muscle cells, fibroblasts, keratinocytes and chondrocytes. The mesenchymal stem cells (MSCs) may be derived from any tissue including bone-marrow, peripheral blood, menstrual fluid, salivary gland, skin and foreskin, synovial fluid, endometrium, dental tissue, adipose tissue and neonatal birth-associated tissues including placenta, umbilical cord, cord blood, amniotic fluid and amniotic membrane. Preferably, the MSCs are derived from bone-marrow or umbilical cord.

In a preferred embodiment, the pre-polymerized solution comprises a cell-line. More preferably, the cell-line is Human Umbilical Vein Endothelial Cells (HUVEC). The multilayered composite can also be obtained using a pre-polymerized solution that comprises 1 to 2% w/v of chitosan, preferably 1 % w/v of chitosan. The pre-polymerized solution of chitosan can additionally comprise at least one compound selected from gelatin, gellam gum, collagen, elastin, cellulose, viable cells, proteins, extracellular vesicles, genetic material or polynucleotides, drugs and polymeric particles. Preferably, the pre-polymerized solution comprising chitosan also comprises endothelial cells and/or mesenchymal stem cells. When chitosan is used in the pre-polymerized solution, the polymerization is obtained by means of a gelling agent. Usually the gelling agent is a solution comprising 2 to 6% w/v, preferably 4% w/v of a gelling compound, such as tripolyphosphate.

The fiber layer may be deposited using any method known in the art. Preferably, the fiber is deposited using electrospinning, melt-spinning and/or solution blow spinning. More preferably, the depositing of the fiber layer is performed by means of solution blow spinning. In solution blow spinning, the spraying apparatus consists of an inner and a concentric outer nozzle. A syringe pump injects and controls the polymer flow through the inner nozzle while compressed air at a certain pressure (PI) flows through the outer nozzle. Because of the nozzle geometry, a region of low pressure around the inner nozzle (P2) is created and a cone is formed by the polymer solution [35]. In a preferred embodiment, the fiber layer comprises poly(ester carbonate urethane)urea (PECUU), polycarbonate urethane)urea, (PCUU), PGA, poly(lactic-co-glycolic acid)(PLGA), poly(lactic acid) (PLA), polyethylene glycol (PEG), fibroin, gelatin and/or polycaprolactone. Preferably, polycaprolactone. The fibers may have a diameter ranging from 10 μιη to 100 nm. Preferably, from 10 μπι ΐο 300 urn. The presence of spun of fibers, such as polycaprolactone fibers, provides structural support to the multilayered composite. The fibers are applied by solution blow spinning (SBS) which is a technique known in the art. When polycaprolactone is used for fiber formation, it can have an average molecular weight Mw from 5000 to 110000 Da, preferably from 50000 to 95000 Da and even more preferably from 60000 to 85000 Da. The solution of the polycaprolactone used for the SBS can comprise from 1% to 30% w/v of polycaprolactone, preferably 1% to 20% w/v and even more preferably from 7% to 15% w/v. Suitable solvents for the polymer solution to be applied by SBS solution includes all solvents in which the polymers are soluble. For example, when polycaprolactone fibers have to be applied, the polycaprolactone can be dissolved in a mixture of acetone and chloroform. Suitable amount can be for example acetone/chloroform 20%/80%. However, other solvents and mixture of solvents in variable amounts can also be used.

The injection rate of the solution in the SBS step can be, for example, from 40μΐνιηίη to 350 μΐνιηιη, preferably from 80 μΐνηιίη to 250 μΐνηιίη and even more preferably from 120 μΐνηιίη to 200 μΐνιηιη. However, also different injection rates can be applied. The air pressure in the SBS step can be, for example, from 10 psi to 120 psi, preferably, from 20 psi to 100 psi, more preferably from 30 psi to 80 psi and even more preferably from 40 psi to 60 psi. However, also different values of air pressure can be applied.

In a preferred embodiment, the layers are formed by depositing fibers in opposite angles to form a mesh. In this embodiment, the template is moved or rotated in one direction to create an arrangement of aligned fibers with a defined angle with respect to a template axis. Then the template is moved or rotated in the opposite direction to create a second arrangement of aligned fibers with an opposite angle. Angled arrangement of fibers is obtained by targeting the output of the fibers source towards the template at the same corresponding angle with respect to the chosen axis of the template. This creates two arrangements of aligned fibers with opposite angles which interweave to form a mesh. In the present examples, a rod is spun while rotating clockwise at 42 rpm, targeting the output of the fiber source at a defined angle with respect to the circumferential axis of the rod. Then the rod is spun while rotating anti-clockwise at 42 rpm, targeting the output of the fiber source at the opposite angle. This exemplifies the type of movement and configuration which can be done to obtain arrangement of fibers which are of opposite orientation in a cylindrical construct. In a preferred embodiment, the fiber layer is deposited in opposite angle and/or phases, and with a wave pattern to form a mesh. In this embodiment, the template is moved or rotated in a particular way to create one wave pattern of aligned fibers and then moved or rotated in the opposite way to create a second wave pattern of aligned fibers. This creates two wave patterns with opposite angles and/or phases which interweave to form a mesh. In the present examples, the angle of spun fiber deposition is defined as explained in the previous paragraph, but the rod is spun for 1 s in a clockwise rotation at 42 rpm followed by a 0.5 s anti-clockwise rotation at 42 rpm and then the rod is spun 1 s anti-clockwise rotation at 42 rpm followed by a 0.5 s clockwise rotation at 42 rpm. This exemplifies the type of movement which can be done to obtain two wave patterns which are of opposite angle.

In a preferred embodiment, the naturally occurring fiber angle is 10 to 80° with respect to a template axis and the fiber layer is deposited at -10 to -80° and/or 10 to 80°. As stated earlier, the naturally occurring fiber angle of a tissue can be calculated from physical or experimental evidences, or by using the Holzapfel's model and parameters obtained from mechanical testing of natural tissues. Beside the described and preferred sequences of clockwise/anti-clockwise rotations (1 s/0.5 s), different time ratios of clockwise/anti-clockwise and rotation velocities could derive into appropriated wavy pattern depositions, this would be apparent for the skilled person after reading the present specification. Additionally, vibration or rapid and short movements of the concentric nozzle system of the solution blow spinning while depositing the fibers, could also derived in wavy pattern deposition. Such rapid and short movement can be up and down (vertically) or right and left (horizontally). Furthermore, the polycaprolactone fibre solution can be supplemented with charged molecules or particles, therefore, spun polycaprolactone fibres projected from the nozzle to the construct under fabrication, magnetic or electrical field pulses can generate wavy pattern of deposited fibers. Likewise, atmospheric perturbations between the nozzle and de fibres deposition zone, through electromagnetic pulses, sound pulses, pressure changes could derived in different levels of waviness of the deposited fibres. In some cases, it may be advantageous to add an alginate layer before applying the polymer fibers, such as the polycaprolactone fibers, by solution blow spinning. This is because a layer of alginate between the layer of polymerized gelatin (i.e. the hydrogel) and the polymer fibers may prevent drying of the polymerized gelatin (i.e. the hydrogel). The drying of the hydrogel layer may also depend on the air stream applied during SBS.

In order to easily remove the multilayered composite from the template, a soft hydrogel structure fabricated from a sacrificial material can be used to reduce or eliminate the damaged caused by friction when removing the multilayered composite from the template. This can be particularly advantageous when the first layer comprises encapsulated cells. Therefore, in a preferred embodiment, the template is dipped at least once in a solution comprising alginate or salts thereof or derivatives thereof and then dipped in a solution inducing the polymerization of the alginate or salts thereof or derivatives thereof before steps (a) to (g).

The alginate solution can comprise from 0.025% to 3% w/v of alginate, salts or derivatives thereof. Suitable salts are for example sodium, lithium and potassium. Preferably, the alginate solution comprises 2% of alginate, salts or derivatives thereof. The solution preferably has a viscosity of 6 cPs (centipoises) to 245 cPs. For the viscosity measurements, the prepared solutions were equilibrated between 20 and 25 minutes at 37 °C before measurement. An Anton Paar MCR 301 rheometer equipped with a cone-plate geometry (plate diameter of 50 mm and cone opening angle of 0.5°) was used to investigate the shear rate dependence of the solution viscosity. A shear flow test with shear rate ramp from 10 to 1000 s^"1 was performed at 37°C. Viscosity data showed a shear rate of 100 s ¹.

As a solution suitable to crosslink or polymerize alginate or salts thereof or derivatives thereof, any solution comprising Ca⁺², Ba⁺², Sr⁺², Fe⁺³, Af³ and the like can be used. Some examples include CaCk, CaSC , CaCC>3 etc. Preferably, said solution comprises CaCi₂. Suitable amount of these compounds in the polymerizing solution are from 1% to 20% w/v, preferably from 2% to 10% w/v and even more preferably 4% to 6% w/v.

After dipping in the polymerization solution, the template coated with alginate can be immersed in a cleansing solution, such as a PBS solution, to remove the polymerization reagent. Optionally, the template coated with alginate can be subjected to successive dipping rounds wherein the successive dipping rounds comprise: a) a first round of 2 submersions in the alginate solution followed by dipping in the polymerization solution and, optionally, in the cleansing solution; or

b) 1 , 2, 3 or 4 submersions in the alginate solution followed by dipping in the polymerization solution and, optionally, in the cleansing solution.

Preferably, the dipping upwards-speed when dipping the rod in the alginate solution is 138 mm/s.

In a preferred embodiment, the multilayered composite is preconditioned by stretching it to at least 120 % of its original length. Preferably, 130 % of its original length. In the interest of clarity, the multilayered composite is stretch so that its length increases by at least 20 %, preferably 30 %. The axial length is the vector at which forces are applied to stretch the material.

In a preferred embodiment, the multilayered composite is stretched and then allowed to relax at least 2 times, preferably 5 times. In the present examples, the multilayered composite was stretched up to 130 % of its original length 5 times. As shown in Figure 2, the preconditioning was necessary to obtain a J curve in the stress-strain curves.

In a preferred embodiment, the template is a rod and the multilayered composite is a multilayered hollow tube. Preferably, in this particular embodiment, the polycaprolactone fiber is deposited using solution blow spinning. More preferably, in this particular embodiment, the gelatin is functionalized using a chemical agent which comprises mathcryloyl. Most preferably, the multilayered hollow tube is stretched and allowed to relax at least 5 times. In an alternative embodiment, the template is dipped horizontally and is half submerged into a pre-polymerized solution comprising gelatin and a photo-initiator, wherein the gelatin is chemically functionalized to become reactive to polymerization or cross-linking in the presence of free radicals. The template is rotated to expose the non-submerged section comprising the pre- polymerized solution to a wavelength of light which stimulates the photo-initiator and causes the gelatin to polymerize or cross-link. Simultaneous to this process, fibers are deposited at an equal or opposite angle to a naturally occurring fiber angle, wherein the template is rotated or moved so that the fibers are deposited in a wavy pattern. Afterward the concentric layer is removed through a longitudinal incision, resulting in a multilayered composite sheet. This method is encompassed within the scope of the first aspect of the present invention and can be used to make, for example, skin grafts.

Multilayered composite

In a second aspect, the present invention provides a multilayered composite obtained or obtainable through any of the methods described herein.

In a preferred embodiment, a layer comprising gelatin (i.e. the hydrogel layer) is 1 μπι to 1 mm thick. In a preferred embodiment, a fiber layer is 10 nm to 5 μπι thick. In a preferred embodiment, the mass of a fiber layer is 0.05 mg/cm² to 10 mg/cm².

In a preferred embodiment, the multilayered composite comprises 1 to 100 hydrogel layers, preferably 1 to 20 hydrogel layers.

The multilayered composite may also be described by the following embodiments:

[1] A multilayered composite comprising: (i) at least one hydrogel layer;

(ii) at least one fiber layer, wherein the layer is formed by depositing fibers in a wavy pattern which forms a mesh and the fibers are deposited at an equal or opposite angle to a naturally occurring fiber angle.

[2] The multilayered composite according to clause [1], wherein the mesh is formed by depositing fibers in opposite angles and/or phases.

[3] The multilayered composite according to any one of the preceding clauses, wherein the hydrogel layer comprises functionalized gelatin, wherein the gelatin is chemically functionalized to become reactive to polymerization or cross-linking in the presence of free radicals.

[4] The multilayered composite according to any one of the preceding clauses, wherein the fibers are synthetic polymers, preferably polycaprolactone fibers.

[5] The multilayered composite according to any one of the preceding clauses, wherein the hydrogel layer is 1 μπι to 1 mm thick.

[6] The multilayered composite according to any one of the preceding clauses, wherein the fiber layer is 10 nm to 5 μπι thick.

[7] The multilayered composite according to any one of the preceding clauses, wherein the naturally occurring fiber angle is 10 to 80° with respect to a template axis and the fiber layer is deposited at a -10 to -80° or 10 to 80° angle. [8] The multilayered composite according to any one of the preceding clauses, wherein the mass of the fiber layer is 0.05 mg/cm² to 10 mg/cm².

[9] The multilayered composite according to any one of the preceding clauses, wherein the tissue comprises 1 to 100 hydrogel layers, preferably 1 to 20 hydrogel layers.

[10] The multilayered composite according to any one of the preceding clauses, wherein the hydrogel layer comprises cells. [11] The multilayered composite according to clause [10], wherein the cells are chosen from a list consisting of mesenchymal stem cells, endothelial cells, smooth muscle cells, fibroblasts, keratinocytes and chondrocytes.

Uses

In a third aspect, the present invention provides the use of any one of the multilayered composites of the present invention for the replacement or patching of blood vessels, skin, cartilage, tendons, ligaments, fistulas, stomach, esophagus, intestines, uterine tubes, larynx, urethra, cardiac tissue or nerve guidance conduits. Preferably, for the replacement or patching of blood vessels or skin.

Examples

Example 1: Coronary Artery Vascular Graft

Preparation of methacryloyl gelatin-alginate (GEAL) solution Methacryloyl gelatin-alginate (GEAL) was synthesized following a previously described protocol [36, 37]. Briefly for Methacryloyl gelatin (GELMA) synthesis, a 10% (w/v) bovine gelatin (Bloom 220, Rousselot, Netherlands) solution in PBS lx (pH 7.4) was prepared and maintained under agitation at 60°C. Methacrylic anhydride (Sigma, US) was added drop-wise to a final concentration of 8% (v/v), allowing the functionalization reaction to occurred for 3 hrs. Methacryloyl functionalization was stopped after adding 3 volumes of PBS IX, and latter submitted to 7 days of dialysis (cut-off molecular weight of 8 kDa) to remove the non-reacted methacrylic anhydride. The solution was freeze dried and stored at room temperature for later use. Three stock solutions were prepared. First, the GEAL stock solution was prepared by dissolving freeze dried GELMA in PBS IX at 40°C at a concentration of 20% (w/v). Then, a 2% (w/v) alginate stock solution was prepared by dissolving medium viscosity sodium alginate (A2033, Sigma, USA) in PBS IX under continuous stirring at 60°C. For the preparation of a photoinitiator (PI) stock solution, 2-Hydroxy-4'-(2-hydroxyethoxy)-2-methylpropiophenone (410896, Sigma, USA) was fully dissolved in PBS IX at 85 °C to obtain a concentration of 2% (w/v), maintained at that temperature before used to avoid crystallization of the reagent. The GEAL solution was obtained after mixing the three stock solutions and the volume adjusted using PBS IX to a final concentration of 10% (w/v) of GELMA, 0.5% (w/v) of alginate and 0.2% (w/v) of PI. Sacrificial alginate scaffold

A sacrificial alginate scaffold was deposited coating a plastic mandrel of 2.5 mm in diameter using the CNC machine following a previously described protocol [38]. Briefly, a 2% (w/v) alginate solution was prepared by dissolving medium viscosity sodium alginate (A2033, Sigma, USA) in PBS IX under continuous stirring at 60°C. For the crosslinking of alginate scaffold, CaC12 was dissolved in ddH20 at a concentration of 5% (w/v) and maintained at 4°C The sacrificial alginate scaffold was built after two subsequent dippings in the alginate solution, then submerged during 15 s in the CaC12 solution for crosslinking and finally immersed tree times in PBS for 1 min for cleansing. Deposition of PCL sub-layers

PCL sublayers were fabricated using PCL spun fibers using a combination of solution blow spinning (SBS) system [35, 39] and a dipping-spinning machine [38]. PCL (440744, Sigma- Aldrich, USA) was dissolved in a mixture of 80/20 ratio of chloroform/acetone to reach final PCL concentration of 7% (w/v). The solution blow spinning configuration is illustrated in Fig. 3b. The system comprises an air compressor (Huracan 1520, Indura, Chile) equipped with a pressure regulator adjusted at 60 psi as a source of blow air; additionally, a 10 mL hypodermic syringe is mounted in a syringe pump (NE-4002X, New Era Pump Systems, Inc. NY, USA) to control the injection rate of the PCL solution at 120 μΕ/ηιίη. Both, the compressed air and the flowing PCL solution are connected and converged into a spraying apparatus that consists of a concentric nozzle system with a central flow of PCL solution a peripheral flow of pressurized air (Fig 1). For fiber deposition, the system requires the dipping- spinning machine coupled to a deposition rod that moves downward and upward at a rate of 138 mm/s while spinning the same rod at 42 rpm. This configuration allows a homogeneous fiber deposition along the spinning rod. A complete cycle of spinning down-and-up movement takes 30 sec and the distance between the SBS nozzle and the point of fiber deposition on the rod surface was kept constant at 30 cm.

For the fabrication of the adventitia layer, the SBS nozzle was orientated at 67° degree with respect to the circumferential axis of the graft, while for the media layer the orientation was 21°. This allowed the placement of aligned fibers in a specific orientation. The fibers were deposited while the rod was subjected to down-and-up ward movement. A simultaneous alternated spinning that consisted of cycles of 1 s clockwise rotation at 42 rpm followed by a 0.5 s anti-clockwise rotation at 42 rpm allows the inclusion of waviness at the oriented PCL fibers. This particular approach was adapted in order to increase the waviness of fibers and imitate the natural configuration of collagen fiber [40-42]. To deposit a sequence of PCL fibers in opposite orientation in order to form an angled mesh comprising the PCL sublayer, the whole dipping- spinning device was switched from forward to backward position, and the spinning movement was changed to 1 s anti-clockwise rotation at 42 rpm followed by a 0.5 s clockwise rotation at 42 rpm. Figure 3 shows the difference between the deposition of stretched fibers (or normal deposition of fibers) and wavy fibers.

Deposition of individual GEAL sub-layers

GEAL layers were generated with a CNC machine [38]. Each layer is fabricated through several dippings of a rod previously covered with a PCL sublayer into the GEAL solution. The gelatin- based pre-crosslinked hydrogel solution was kept in a water bath at 30°C to avoid spontaneous gelation at room temperature. Crosslinking was achieved during emersion of the rod by exposing the temporally coated rod (gelatin solution is briefly stabilized before photo-crosslinking by cohesive forces) to UV light at 365 nm wavelength (1.21 W/cm2) (OmniCure® S2000, Excelitas Technologies, USA). The UV source is place at a distance of 2 cm from the rod while the coated mandrel was rotating at 42 rpm and emerging at 138 mm/s upward- speed.

The middle and outer layer, which represent the media and adventitia layer of a natural artery, were fabricated intercalating GEAL and a PCL sub-layers. First, a thin sacrificial alginate scaffold was deposited around the plastic rod to allow a gentle removal of the cylindrical multi- layer construct after fabrication. For the middle PCL sub-layers, the SBS nozzle was oriented at - 21° with respect to the circumferential axis. First, a complete spin-down-upward movement of the rod was performed with the CNC machine in a forward position while blowing PCL fibers for deposition at angle of +21 °. Then a spin-down-upward cycle was performed using the CNC machine in a backward position in order to deposit the fibers at -21 °. Afterward, the formation of a concentric GEAL sub-layer was carried out by dipping the fiber-coated rod into the GEAL solution and kept there for 30 s in order to allow the pre-crosslinking solution to permeate through the fibers of the PCL sub-layer. Subsequently, the previously described process of emersion and UV-crosslinking was applied to stabilize the hydrogel within the "graft sublayer". One "graft sublayer" consists of typically one PCL sublayer and one crosslinked GEAL sublayer. The same general methodology was used for the outer layer fabrication but changing the SBS nozzle angle to +67° and -67°.

Different composition of sublayers and layers were tested in order to determine the best composition for mechanical response. For this reason, the GEAL sublayers can be composed of 1 , 2 or 3 cycles of dipping and UV-crosslinking. Additionally, the complete middle and outer layer can comprise series of 4, 5, 8, or 10 middle or outer graft sublayers respectively. The better performing middle layer was fabricated with 4 series of middle graft sublayer, consisted at the same time of a PCL sublayer and a GEAL sublayer generated after 2 cycles of dipping and UV- crosslinking. The best outer layer was consisted of 5 series of outer graft sublayer, in which the GEAL sublayer required 3 cycles of dipping and UV-crosslinking.

Figure 4 shows different compositions for middle and outer layers. The layer formulation in which its mechanical response most closely approximates the native adventitia layer, hereafter called outer layer, is the one manufactured with 5 series of outer graft sublayer, comprised at the same time of a 1 PCL sublayer and 1 GEAL sublayer formed after 3 cycles of dipping- spinning/crosslinking in GEAL solution. On the other hand, for mimicking the media layer mechanics, the most similar formulation was the one fabricated with 4 series of middle graft sublayer, consisted as well of 1 PCL sublayer and 1 GEAL sublayer formed after 2 cycles of dipping-spinning/crosslinking in GEAL solution, hereafter called middle layer. Inspired by the natural distribution of intercalated collagen/elastin fibers and cells in human arteries, an intercalated configuration of PCL fiber sub-layer and GEAL sublayer deposition was chosen.

Fabrication of a full vascular graft using GEAL layers reinforced with PCL fibers

A full vascular graft based on the reinforced GEAL hydrogel consists of three layers: inner, middle and outer, mimicking the tissue configuration of native coronary arteries (intima, media, and adventitia, respectively). In order to obtain an unscathed inner cellularized layer after removal of the vascular graft from the plastic rod, a thin coating of sacrificial alginate scaffold was first generated around the plastic rod. This allowed a gentle removal of the supporting rod from the inside of the vascular graft. For the inner layer, a GEAL sub-layer was fabricated after 9 cycles of dippings and UV-crosslinking using the alginate-coated rod and the CNC machine. On top of the inner layer, an optimized middle layer was manufactured as described above. Subsequently, the concentric optimized outer layer was fabricated around the middle layer. Finally, the plastic rod and alginate coat were removed mechanically.

Tensile test

Uniaxial tensile test for the middle layer and outer layer were performed in a Texture analyzer (Stable Micro Systems, TA.XT.plus, Surrey, UK). The axial force was measured with a 5 N load cell. After fabrication, rectangular sections with circumferential and longitudinal orientation were cut and maintained at 37° in PBS IX (pH 7.4). For each layer, 3 samples were tested for circumferential tensile testing and 3 for longitudinal testing. Sample thickness and width were measured for each sample using a micrometer caliper with 0.01 mm of accuracy. For mechanical analysis, the sample length was considered as the distance between clamps in the texture analyzer after positioning the sample at the beginning of the test. Before the testing was carried out, five loading and unloading cycles at a constant rate of 10 mm/s were applied as the preconditioning step. The preconditioning loading/unloading cycles for the longitudinal test of the outer layer included a maximum strain of 13%, while for the circumferential test; the maximum strain was 30%. In the case of the middle layer, maximum strain of preconditioning was 35% and 30% for the circumferential and longitudinal test respectively. Axial testing for both, circumferential and longitudinal samples, were performed at a constant rate of 10 mm/s.

Uniaxial tensile test for the full vascular grafts were performed in a universal testing machine (Instron 3342, Norwood, MA, USA). The axial force was measured with a 10 N load cell. The rectangular samples were cut in longitudinal and circumferential directions. The samples were maintained and tested while being permanently submerged in phosphate buffered saline (PBS) IX at a temperature of 37° ± 0.5°C. For each vascular graft, 5 samples were tested. Sample thickness, width and length were measured with an optical extensometer with 0.001 mm of precision. Before carrying out the tensile testing, five loading/unloading cycles to a strain of 30% at a constant rate of 10 mm/s was used as preconditioning step. The axial circumferential and longitudinal testing of samples were performed at a constant rate of 1 mm/s. In order to test the resistance to circumferential deformation, the full vascular graft was cut in the circumferential direction and was subjected to 20 loading/unloading cycles of circumferential stress at maximum strain of 30% and constant rate of 10 mm/s.

Stress-strain curves for all tests were derived from axial load and clamps displacement recorded along the test. The stress was computed as F/A, where the F is the axial load with a precision of 0.01N and A is the initial cross-sectional area. The strain was computed as 100*L/L0, with L and L0 as the current length and initial sample length, respectively.

In Figure 5, tensile test in circumferential and longitudinal directions were performed to the optimized middle and outer layers. Each layer showed an anisotropic and nonlinear mechanical response resembling the media and adventitia layers of human coronary arteries, respectively. The outer layer showed a stiffer behavior in the longitudinal direction compared to the circumferential direction (Fig. 5a, 5b), whereas the middle layer is stiffer at the circumferential direction (Fig. 5c, 5d). An opposite response is observed between the middle and outer layer in each direction. At the longitudinal direction, the middle layer is more compliant (Fig. 5c); whereas at the circumferential direction middle layer tends to be slightly stiffer than the outer layer (Fig. 5b and 5d).

In Figure 6, tensile test in circumferential and longitudinal directions were performed to the complete vascular graft, and additionally a cyclic load and unload tensile test was performed in the circumferential direction to evaluate the fatigue of the material. In Figure 6a and 6b, the full vascular graft maintained the nonlinear response and a clear anisotropy. In addition, a similar mechanical behavior to human coronary arteries was achieved in both tensile testing directions, almost always inside the range of healthy natural coronary arteries. In Figure 6c, to verify whether the mechanical properties of the vascular graft change upon the application of repeated loading and unloading cycles, simulating the pulsatile flow under in vivo physiological conditions, 20 repetition of circumferential stress were performed to a strip taken from the fabricated vascular graft. A linear response is shown in the first cycle, which would correspond to the pre-condition stretching step. However, after the first cycle a non-linear and anisotropic behavior is observed, converging to a J-shaped stress-strain curve with small hysteriesis.

Pressurization test

A pressurization test was used to study the response of the full vascular graft under simulated conditions of human in vivo loading and pressure conditions. The test was performed in a customized set up using a universal testing machine (Instron 3342, Norwood, MA, USA), adapted with a plastic transparent chamber filled with PBS IX and with controlled temperature at 37° ± 0.5°C. The internal pressure was applied using an auxiliary line of PBS at 37°C connected to the internal graft lumen. The pressure was measured at the entrance of the chamber with a pressure transducer, whereas the graft diameter with an optical extensometer. Five samples with an average length of 5 mm were tested. Before testing, five loading/unloading cycles of longitudinal strain at 30% and a constant rate of 10 mm/s were carried out as preconditioning step. An additional preconditioning step in the circumferential direction was performed using 5 cycles of pressurization from 0 to 200 mmHg. Pressurization test of the full vascular grafts were subjected under three different constant axial strains of 10%, 20% and 25%.

The compliance value of full vascular grafts (%C) was computed from the experimental data at three pressure ranges (50-90, 80-120, 110-150 mmHg), according to standard ISO 7198 (ANSI/AAMI/ 2010) and using following equation: 2 - 1 '

where PI and P2 correspond to the lower and higher range of pressure values in mmHg, and RP1 and RP2 are the external radiuses generated at those pressures respectively. Figures 7a, 7b, and 7c shows that a J-shape response was obtained at the diameter-pressure curves and was not altered by higher axial pre-stretching deformation after grafts installation and before pressurized testing. Compared with human coronary arteries, vascular grafts showed a more compliant response at lower ranges of luminal pressures. In the approximated in vivo axial stretching of natural blood vessels (10 %), vascular grafts showed a greater increase in external diameter under luminal pressurization compared with human coronary arteries (Fig 7. a). At 20% and 25% axial elongation during pressurized mechanical testing, similar values of nominal diameter change to human coronary arteries are observed (Fig 7.b and Fig7.c).

In Table 1 , Table 2 and Table 3 it can be seen the compliance values for vascular grafts with 10 and 20% of axial stretching condition have no statistical difference with human coronary arteries. These results highlight the statistically similar compliant response that vascular grafts have at pressures close to physiologic ranges (80-120 mmHg), making them specially prepared to response properly in a theoretical bypass situation. At higher values of axial elongation (30 %) compliance values of human coronary arteries and vascular grafts were statistically different (Table 3), hence axial stretching conditions during bypass surgery need to be considered.

IBllH

Pressure range

50-30 30-120 110-150

ascular graft 24.0 + 4. S 12.9 + 2.60 7.5 + OSS

Coronary arte ry 7.1 ± 0 17 " 2.7 + 037 ' 1.2 ± 0.65"

¾tatis 3iis^'i i fe n e ίρ ί i S ¾ com :ρ:κ-¾< ό wl t v ssvTU i ar graft

"Statistics: difference ·:ρ<ϊ J.325^'s tampered viitti yascuiar graft

Viability and Proliferation tests

Human Umbilical Cord Cells (HUVEC) (ATCC® CRL1730™) were cultured and expanded in culture medium (high glucose Dulbecco's Modified Eagle's medium (DMEM) (16000-044, Gibco, USA) supplemented with 10%(v/v) fetal bovine serum (FBS) (16000-044, Gibco, USA), 2 mM glutamine (25030-081, Gibco, USA) and 1 % (v/v) penicillin-streptomycin (15140-122, Gibco, USA)) and incubated at 37°C, 5% C02 and 96% of humidity. HUVECs were mixed in the GEAL solution at a concentration of 10 million cells mL-1. Vascular grafts were fabricated as mentioned before using a mixture of GEAL solution and HUVECs in order to encapsulate the cells within the GEAL sublayers.

Cell proliferation tests were performed using a 5mm cylindrical section of the full vascular grafts using the WST-1 Cell Proliferation Colorimetric Assay Kit (K302, Biovision, USA). Briefly, this assay quantifies the metabolic cleavage of WST-1 to generate formazan by cellular mitochondrial dehydrogenases. The proliferation was measured at day 1 and day 7 after cell culture of the vascular graft sections in culture medium supplemented with amphotericin b (15290-026, Gibco, USA). Cylindrical samples were afterward washed in PBS before including them into 200 μL· of culture medium and 20 μL· of WST reagent, and incubated for 2,5 hrs. A standard curve was done to estimate the number of cells in the vascular grafts. For this, 5000, 10000, 20000 and 30000 cells were seeded and incubated with 200 μL· of culture medium and 20 μΕ of WST reagent for 2,5 hrs. Number of metabolically active cells in the vascular graft was obtained interpolating the absorbance values in the standard curve. Cell density was calculated dividing the number of active cells by the vascular graft volume («30 mm3).

In Figure 8b, it can be seen a cell proliferation assay that was performed to evaluate the cell damage generated during the manufacturing process. The active cell density on day 1 was low in comparison with the cell density previously present in the GEAL solution or compared to the cell density quantified in the stained section. However, a great increase of cells activity is observed at day 7, with significance difference with day 1. The same day of fabrication, vascular grafts were embedded in O.C.T. Compound (Tissue-Tek, USA) and sectioned at 14 μπι in transversal cuts using a cryostat (Microm, HM525, Walldorf, Germany). For the cell staining, the samples were incubated in Hoechst 33342 solution (Thermo Scientific, USA) following the provider's protocol. Transversal cuts were visualized using a fluorescent microscope (CKX41, Olympus, USA).

Figure 8a shows the stained section showing a high cell density (22.000 cells/mm3). Vascular grafts also showed homogeneous and concentric distribution of cells throughout the thickness of the graft wall and aligned distribution of cells following the concentric positioning of fabricated sub-layer.

Example 2: Skin graft GEAL solution, sacrificial alginate and deposition of GEAL and PCL fiber sublayers

The same protocols mentioned in Example 1 were used to produce the GEAL solution and the sacrificial alginate scaffold and the deposition of GEAL sub-layers and PCL fibers sublayers to produce skin grafts. The fiber deposition orientation was determined using the Holzapfel model as mentioned before. For the fabrication of the skin graft, the SBS nozzle was orientated at 30° degree with respect to the circumferential axis of the graft and the same protocol mentioned above to obtain waviness in fibers and preconditioning of the graft was done. Fabrication of a full skin graft using GEAL layers reinforced with PCL fibers Skin grafts were fabricated intercalating GEAL and a PCL sub-layers. First, a thin sacrificial alginate scaffold was deposited around the plastic rod to allow a gentle removal of the cylindrical multi-layer construct after fabrication. The SBS nozzle was oriented at -30° with respect to the circumferential axis. First, a complete spin-down-upward movement of the rod was performed with the CNC machine in a forward position while blowing PCL fibers for deposition at angle of +30°. Then a spin-down-upward cycle was performed using the CNC machine in a backward position in order to deposit the fibers at +30°. Afterward, the formation of a concentric GEAL sub-layer was carried out by dipping the fiber-coated rod into the GEAL solution and kept there for 30 s in order to allow the pre-crosslinking solution to permeate through the fibers of the PCL sub-layer. Subsequently, the previously described process of emersion and UV-crosslinking was applied to stabilize the hydrogel within the "graft sublayer". The procedure of dipping and emersion with UV-crosslinking was repeated 3 times in order to obtain a crosslinked GEAL sublayer. One "graft sublayer" consists of typically one PCL sublayer and one crosslinked GEAL sublayer. To produce the final skin graft 3 or 15 grafts sublayers were deposit over the rod.

Tensile Test

Uniaxial tensile test for the middle layer and outer layer were performed in a Texture analyzer (Stable Micro Systems, TA.XT.plus, Surrey, UK). The axial force was measured with a 5 N load cell. After fabrication, rectangular sections with longitudinal orientation were cut and maintained at 37° in PBS IX (pH 7.4). Sample thickness and width were measured for each sample using a micrometer caliper with 0.01 mm of accuracy. For mechanical analysis, the sample length was considered as the distance between clamps in the texture analyzer after positioning the sample at the beginning of the test. Before the testing was carried out, five loading and unloading cycles at a constant rate of 10 mm/s were applied as the preconditioning step. The preconditioning loading/unloading cycles for the longitudinal test of the skin graft was 30% of maximum strain. Stress-stain curves were derived as mentioned above.

Figure 9 shows the comparison between native skin and our skin graft. A J-shaped response similar to the native skin is observed in the skin graft stress stain curve for both 3 and 15 cycles.

Bibliography

1. Catto, V., et al., Vascular tissue engineering: recent advances in small diameter blood vessel regeneration. ISRN Vascular Medicine, 2014. 2014.

2. Melchiorri, A.J., N. Hibino, and J.P. Fisher, Strategies and techniques to enhance the in situ endothelialization of small-diameter biodegradable polymeric vascular grafts. Tissue Engineering Part B: Reviews, 2013. 19(4): p. 292-307.

3. Liu, J., et al., Properties of single electrospun poly (diol citrate)-collagen-proteoglycan nano fibers for arterial repair and in applications requiring viscoelasticity. Journal of biomaterials applications, 2014. 28(5): p. 729-738.

4. Tang, C, et al., Platelet and endothelial adhesion on fiuorosurfactant polymers designed for vascular graft modification. Journal of Biomedical Materials Research Part A, 2009. 88(2): p. 348-358. 5. Liu, H., et al., Improved hemocompatibility and endothelialization of vascular grafts by covalent immobilization of sulfated silk fibroin on poly (lactic-co-glycolic acid) scaffolds. Biomacromolecules, 2011. 12(8): p. 2914-2924. 6. Klinkert, P., et al., Saphenous vein versus PTFE for above-knee femoropopliteal bypass. A review of the literature. European Journal of Vascular and Endovascular Surgery, 2004. 27(4): p. 357-362.

7. Yue, X., et al., Smooth muscle cell seeding in biodegradable grafts in rats: a new method to enhance the process of arterial wall regeneration. Surgery, 1988. 103(2): p. 206-212.

8. Niklason, L., et al., Functional arteries grown in vitro. Science, 1999. 284(5413): p. 489-493.

9. Shum-Tim, D., et al., Tissue engineering of autologous aorta using a new biodegradable polymer. The Annals of thoracic surgery, 1999. 68(6): p. 2298-2304.

10. Dardii, H. and N. Engkwood, Evaluation of glutaraldehyde-tanned human umbilical cord vein as a vascular prosthesis for bypass to the popliteal, tibial, and peroneal arteries. Age (mean), 1978. 67(64): p. 74.

11. Dardik, H., et al., Biodegradation and aneurysm formation in umbilical vein grafts. Observations and a realistic strategy. Annals of surgery, 1984. 199(1): p. 61. 12. Neufang, A., et al., Sequential femorodistal composite bypass with second generation glutaraldehyde stabilized human umbilical vein (HUV). European journal of vascular and endovascular surgery, 2005. 30(2): p. 176-183. 13. Borschel, G.H., et al., Tissue engineering of recellularized small-diameter vascular grafts. Tissue Engineering, 2005. 11(5-6): p. 778-786.

14. Cho, S.-W., et al., Evidence for in vivo growth potential and vascular remodeling of tissue- engineered artery. Tissue Engineering Part A, 2008. 15(4): p. 901-912.

15. Brennan, J. A., J.H. Arrizabalaga, and M.U. Nollert, Development of a Small Diameter Vascular Graft Using the Human Amniotic Membrane. Cardiovascular Engineering and Technology, 2014. 5(1): p. 96-109. 16. Cho, S.-W., et al., Small-diameter blood vessels engineered with bone marrow-derived cells. Annals of surgery, 2005. 241(3): p. 506.

17. Dahl, S.L., et al., DeceUularized native and engineered arterial scaffolds for transplantation. Cell transplantation, 2003. 12(6): p. 659-666.

18. Dahl, S.L., et al., Readily available tissue-engineered vascular grafts. Science translational medicine, 2011. 3(68): p. 68ra9-68ra9.

19. Hirai, J., et al., Highly oriented, tubular hybrid vascular tissue for a low pressure circulatory system. ASAIO journal, 1994. 40(3): p. M383-M388. 20. Lovett, M., et al., Silk fibroin microtubes for blood vessel engineering. Biomaterials, 2007. 28(35): p. 5271-5279.

21. Sadr, N., et al., SAM-based cell transfer to photopatterned hydrogels for microengineering vascular-like structures. Biomaterials, 2011. 32(30): p. 7479-7490.

22. L'Heureux, N., et al., Human tissue-engineered blood vessels for adult arterial revascularization. Nature medicine, 2006. 12(3): p. 361-365.

23. L'Heureux, N., et al., A completely biological tissue-engineered human blood vessel. The FASEB Journal, 1998. 12(1): p. 47-56.

24. Hu, J. -J., et al., Construction and characterization of an electrospun tubular scaffold for small-diameter tissue-engineered vascular grafts: a scaffold membrane approach, journal of the mechanical behavior of biomedical materials, 2012. 13: p. 140-155.

25. Zhou, M., et al., Development and in vivo evaluation of small-diameter vascular grafts engineered by outgrowth endothelial cells and electrospun chitosan/poly (ε-caprolactone) nanofibrous scaffolds. Tissue Engineering Part A, 2013. 20(1-2): p. 79-91.

26. Nieponice, A., et al., In vivo assessment of a tissue-engineered vascular graft combining a biodegradable elastomeric scaffold and muscle-derived stem cells in a rat model. Tissue Engineering Part A, 2010. 16(4): p. 1215-1223. 27. Blaker, J.J., et al. Highly porous nanofibres and their formation into macroporous structures by solution blow spinning: Applications in tissue regeneration, in ABSTRACTS OF PAPERS OF THE AMERICAN CHEMICAL SOCIETY. 2014. AMER CHEMICAL SOC 1155 16TH ST, NW, WASHINGTON, DC 20036 USA.

28. Barry, F.P. and J.M. Murphy, Mesenchymal stem cells: clinical applications and biological characterization. The international journal of biochemistry & cell biology, 2004. 36(4): p. 568- 584. 29. Wang, S., X. Qu, and R.C. Zhao, Clinical applications of mesenchymal stem cells. J Hematol Oncol, 2012. 5(1): p. 19.

30. Hashi, C.K., et al., Antithrombogenic property of bone marrow mesenchymal stem cells in nanofibrous vascular grafts. Proceedings of the National Academy of Sciences, 2007. 104(29): p. 11915-11920.

31. Mirza, A., et al., Undifferentiated mesenchymal stem cells seeded on a vascular prosthesis contribute to the restoration of a physiologic vascular wall. Journal of vascular surgery, 2008. 47(6): p. 1313-1321.

32. Olausson, M., et al., Transplantation of an allogeneic vein bioengineered with autologous stem cells: a proof-of-concept study. The Lancet, 2012. 380(9838): p. 230-237. 33. Holzapfel, 2000. Holzapfel GA, Gasser TC, and Ogden RW. A new constitutive framework for arterial wall mechanics and a comparative study of material models. J Elasticity 61 : 1-48, 2000. 34. Holzapfel, 2005. Holzapfel, G. A., Sommer, G., Gasser, C. T., & Regitnig, P. (2005). Determination of layer-specific mechanical properties of human coronary arteries with nonatherosclerotic intimal thickening and related constitutive modeling. American Journal of Physiology-Heart and Circulatory Physiology, 289(5), H2048-H2058. 35. Medeiros et al, 2009. Medeiros, E. S., Glenn, G. M, Klamczynski, A. P., Orts, W. J., & Mattoso, L. H. (2009). Solution blow spinning: A new method to produce micro-and nanofibers from polymer solutions. Journal of Applied Polymer Science, 113(4), 2322-2330.

36. Nichol et al, 2010. Nichol, J. W., Koshy, S. T., Bae, H., Hwang, C. M., Yamanlar, S., & Khademhosseini, A. (2010). Cell-laden microengineered gelatin methacrylate hydrogels.

Biomaterials, 31(21), 5536-5544.

37. Van den Bulcke et al., 2000. Van Den Bulcke, A. I., Bogdanov, B., De Rooze, N., Schacht, E. H., Cornelissen, M., & Berghmans, H. (2000). Structural and rheological properties of methacrylamide modified gelatin hydrogels. Biomacromolecules, 1(1), 31-38.

38. Wilkens et al, 2016. Wilkens, C. A., Rivet, C. J., Akentjew, T.L., Alverio,. J. , Khoury, M., Acevedo, J.P. (2016). Layer-by-Layer Approach for a Uniformed Fabrication of a Cell Patterned Vessel-Like Construct. Biofabrication, 9(1):015001. 39. Medeiros et al., 2016. Medeiros, E. L. G., Braz, A. L., Porto, I. J., Menner, A., Bismarck, A., Boccaccini, A. R., ... & Blaker, J. J. (2016). Porous Bioactive Nanofibers via Cryogenic Solution Blow Spinning and Their Formation into 3D Macroporous Scaffolds. ACS Biomaterials Science & Engineering, 2(9), 1442-1449.

40. Chen et al, 2011. Chen, H., Liu, Y., Slipchenko, M. N., Zhao, X., Cheng, J. X., & Kassab, G. S. (2011). The layered structure of coronary adventitia under mechanical load. Biophysical journal, 101(11), 2555-2562. 41. Rezakhaniha et al., 2012. Rezakhaniha, R., Agianniotis, A., Schrauwen, J. T. C, Griffa, A., Sage, D., Bouten, C. V. C, ... & Stergiopulos, N. (2012). Experimental investigation of collagen waviness and orientation in the arterial adventitia using confocal laser scanning microscopy. Biomechanics and modeling in mechanobiology, 11(3-4), 461 -473. 42. Chen et al., 2013. Chen, H., Slipchenko, M. N., Liu, Y., Zhao, X., Cheng, J. X., Lanir, Y., & Kassab, G. S. (2013). Biaxial deformation of collagen and elastin fibers in coronary adventitia. Journal of Applied Physiology, 115(11), 1683-1693.

Claims

1. A method for producing a multilayered composite sheet comprising the steps of:

(a) dipping a template into a pre-polymerized solution comprising gelatin and a photo -initiator and alginate or salts thereof, wherein the gelatin in the pre- polymerized solution is chemically functionalized to become reactive to polymerization or cross-linking in the presence of free radicals, and wherein the amount of functionalized gelatin in the pre-polymerized solution is in the range of 1-20% w/v;

(b) exposing the pre-polymerized solution attached to the template to a wavelength of light which stimulates the photo-initiator and causes the gelatin to polymerize or cross-link;

(c) optionally repeating steps (a) and (b) to obtain the desired number of layers;

(d) depositing a fiber layer on the template at an equal or opposite angle to a naturally occurring fiber angle, wherein the template is rotated clockwise and then anti-clockwise so that the fibers are deposited in a wavy pattern, and wherein the fiber layer comprises polycaprolactone;

(e) optionally repeating step (d) to obtain the desired number of layers;

(f) optionally repeating steps (a) to (e) to obtain the desired number of layers; and

(g) optionally preconditioning the resultant multilayered composite sheet by stretching and relaxing it.

2. The method according to claim 1, wherein the multilayered composition sheet is a skin graft.

3. The method according to claim 1 or 2, wherein the gelatin of step (a) is functionalized using a chemical agent which provides methacryloyl, methacrylamide, acrilamide and/or acryloyl functionalization at the amino acid side chain of polymers.

4. The method according to any one of the preceding claims, wherein the pre-polymerized solution comprises viable cells, proteins, extracellular vesicles, genetic material or polynucleotides, drugs and/or polymeric particles.

5. The method according to claim 4, wherein the viable cells are chosen from a list consisting of mesenchymal stem cells, endothelial cells, smooth muscle cells, fibroblasts, keratinocytes and chondrocytes.

6. The method according to any one of the preceding claims, wherein the depositing of the fiber layer is performed by means of solution blow spinning.

7. The method according to any one of the preceding claims, wherein the fiber layer is deposited in opposite angle and/or phases to form a mesh.

8. The method according to any one of the preceding claims, wherein the naturally occurring fiber angle is 10 to 80° with respect to a template axis and the fiber layer is deposited at a -10 to -80° and 10 to 80° angle.

9. The method according to any one of the preceding claims, wherein before steps (a) to (g): the template is dipped at least once in a solution comprising alginate or salts thereof; and then dipped in a solution inducing the polymerization of the alginate or salts thereof or derivatives thereof.

10. The method according to any one of the preceding claims, wherein the composite is stretched to at least 120 % of its original length in step (g), preferably 130 %.

11. The method according to any one of the preceding claims, wherein the composite is stretched and relaxed in step (g) at least 2 times, preferably 5 times.

12. A multilayered composite obtained or obtainable through a method according to any one of claims 1-11.

13. Use of the multilayered composite according to claim 12 as dependent on claims 1 , and 3 to 11 , for the replacement or patching of blood vessels, skin, cartilage, tendons, ligaments, cardiac tissue, stomach, esophagus, intestines, uterine tubes, larynx, urethra or nerve guidance conduits.