implant

Abstract

This invention relates to implants. More specifically, this invention relates to implants comprising a matrix material, and a method for preparing implants comprising a matrix material. This invention also relates to coiled implants.

Description

[0001] This application is a divisional application of the application filed on July 5, 2013, with application number 201380035760.X and invention title "Implant". Technical Field

[0002] This invention relates to implants comprising a matrix material, and to a method for preparing implants comprising a matrix material. The invention also relates to coiled implants. Background of the Invention

[0003] Since the early 1990s, a relatively new field in medicine has been regenerative medicine. Regenerative medicine is the process of creating living, functional tissue to repair, replace, or restore the structure and function of tissues or organs lost due to age, disease, injury, or congenital defects. This field utilizes novel methods, including (stem) cell therapy, the development of medical devices, and tissue engineering.

[0004] In recent years, continuous improvements in our healthcare have led to dramatic demographic changes, such as an increase in the average age of the population. These demographic changes are causing an increase in the prevalence of age-related diseases such as cardiovascular disease. Many of these diseases are caused by the loss or dysfunction of specific cell types in the body, resulting in permanent damage to tissues and organs.

[0005] Cardiovascular disease is one of the leading causes of death worldwide. One approach to treating at least some of these diseases is through tissue engineering. Tissue engineering can be used to replace cardiovascular tissues such as arteries and heart valves. Current cardiovascular replacements face risks due to aggregation, infection, degradation, and lack of growth potential. Tissue engineering uses the patient's own cells and biodegradable polymer scaffolds to enable the autologous tissue to grow, adapt, and repair. To ensure proper cell and tissue growth, the scaffold must be highly porous and match the mechanical properties of the tissue. Electrospinning is a technique that uses a high-voltage electrostatic field to produce polymer nanofibers. It produces highly porous materials composed of nanofibers similar to the extracellular matrix of tissue. Tissue engineering can be used, for example, for coronary artery bypass grafts, heart valve replacements, and arteriovenous bypasses for dialysis patients.

[0006] In the surgical field, minimally invasive surgery is preferred. However, tissue-engineered constructs are often implanted only via conventional surgery because the constructs cannot be compressed to a sufficiently small size to facilitate minimally invasive procedures. Currently, some artificial heart valves can be rolled up to a diameter of 18 French (6 mm), allowing implantation via small peripheral incisions (e.g., via the femoral artery or jugular vein). However, many older patients requiring valve replacement also suffer from stenotic and consequently narrowed arteries, currently excluding them from the highly preferred minimally invasive procedure. Reducing the rollable diameter by 1 or 2 French has already meant a significant increase in the number of patients who can be treated.

[0007] Tissue engineering techniques consist of constructing alternatives (such as biological substitutes) for diseased tissues. Tissue engineering uses natural or polymeric scaffolds, which provide mechanical support and promote the regeneration of cells lost due to trauma or disease. Scaffolds are temporary structures used to support materials (such as tissue) during their recovery.

[0008] Polymer scaffolds can be constructed from biocompatible, non-toxic polymers. The choice of polymers and techniques used to fabricate scaffolds affects the mechanical properties exhibited by the scaffolds.

[0009] In the publication by Bouten et al., Advanced Drug Delivery Reviews, 2011, Vol. 63, pp. 221-241, synthetic polymers have been shown to be good substrates for valve and vascular tissue engineering. For cardiac tissue engineering, the most commonly used biodegradable synthetic scaffold materials are polyglycolic acid (PGA), polylactic acid (PLA), polyhydroxybutyrate (PHB), ε-polycaprolactone (PCL), or copolymers thereof. No functional implants using these synthetic scaffold materials have been disclosed.

[0010] In a publication by Dankers et al., Nature Materials, 2005, Vol. 4, pp. 568-574, solution-cast polymer films containing the 2-ureido-4[1H]-pyrimidinone (Upy) polymer were shown to be non-toxic when studied in vivo. However, the use of the UPy polymer as a cardiovascular implant stent was not shown.

[0011] To obtain tissue-engineered constructs, scaffolds can be pre-seeded in vitro with appropriate cells prior to implantation. In most cases, scaffold degradation should occur slowly and steadily as new tissue forms and remodels, leaving only new healthy tissue behind. Degradation refers to the breakdown of materials into smaller components, such as chemical compounds and / or elements that can be eliminated from the body, for example, through excretion in urine.

[0012] The disadvantage of growing tissue constructs in vitro is that the entire process, including growth and implantation, must be performed aseptically, making it an expensive and labor-intensive procedure. Furthermore, the regulatory guidelines for living tissue are complex, leading to lengthy and costly product approval processes.

[0013] Another option is to seed the implant with cells prior to implantation. This method requires collecting cells from the subject to receive the implant, optionally growing the cells in vitro, seeding the cells in the construct, and then implanting them. This method has the same negative impacts as the previously described methods. Brief description of the attached figures

[0014] Figure 1 This shows the natural fibrous orientation of the heart valve leaflets (Sauren (1981)).

[0015] Figure 2 The orientation of spiral fibers in blood vessels is shown (Holzapfel, J. Bast., 2000).

[0016] Figure 3 The results of a systemic test of the PCL-bisurea valve at 120 / 80 mmHg are shown, demonstrating its stability over 20 hours.

[0017] Figure 4 Images of the PCL-double urea valve are shown at 20 hours of valve testing (R11020), at the start of the systemic condition (two top images), and 20 hours later (two bottom images).

[0018] Figure 5 The results of a valve test on a PCL valve under systemic conditions at 50 / 25 mmHg are shown, indicating a decline in performance over several hours.

[0019] Figure 6 Images of the PCL valve are shown at 20 hours of valve testing (R11020), at the start of 50 / 25 mmHg (two top images) and 20 hours later (two bottom images).

[0020] Figure 7 A photograph (a) and a schematic overview (b) show the valve test setup.

[0021] Figure 8 The results of uniaxial tensile tests of PCL and PCL diurea electrospun scaffolds along (Figure 1) and perpendicular to (Figure 2) the main fiber direction are shown.

[0022] Figure 9A The results of uniaxial fatigue tests (10%, 2Hz) on PCL and PCL diurea are shown. Figure 9BThe results of pipe fatigue tests using PCL bisurea and PCL UPy implants are shown.

[0023] Figure 10 The results of a curling test on electrospun heart valve implants are shown.

[0024] Figure 11 This shows different curling states of a 3D electrospun heart valve.

[0025] Figure 12 The stent is shown after different stages of the curling of a 3D electrospun heart valve.

[0026] Figure 13 Two images showing DAPI staining of PCL-diurea implants in a sheep model are displayed.

[0027] Figure 14 Two SEM images of PCL and PCL-diurea matrix are shown. Summary of the Invention

[0028] The purpose of this invention is to provide an implant that can be implanted via minimally invasive surgery.

[0029] The purpose of this invention is to provide an implant for regenerating tissue in vivo.

[0030] A further object of the present invention is to overcome one or more of the disadvantages associated with the prior art mentioned above.

[0031] One or more of the above-mentioned objectives have been achieved through this invention. Surprisingly, the inventors have discovered that the above-mentioned objectives are achieved by implants comprising one or more supramolecular compounds, wherein the matrix material comprises a fibrous network and has a porosity of at least 60%, preferably between 70% and 90%.

[0032] The invention will be described in more detail below with reference to the accompanying drawings, in which:

[0033] Figure 1 This shows the natural fibrous orientation of the heart valve leaflets (Sauren (1981)).

[0034] Figure 2 The orientation of spiral fibers in blood vessels is shown (Holzapfel, J. Bast., 2000).

[0035] Figure 3 The results of a systemic test of the PCL-bisurea valve at 120 / 80 mmHg are shown, demonstrating its stability over 20 hours.

[0036] Figure 4Images of the PCL-double urea valve are shown at 20 hours of valve testing (R11020), at the start of the systemic condition (two top images), and 20 hours later (two bottom images).

[0037] Figure 5 The results of a valve test on a PCL valve under systemic conditions at 50 / 25 mmHg are shown, indicating a decline in performance over several hours.

[0038] Figure 6 Images of the PCL valve are shown at 20 hours of valve testing (R11020), at the start of 50 / 25 mmHg (two top images) and 20 hours later (two bottom images).

[0039] Figure 7 A photograph (a) and a schematic overview (b) show the valve test setup.

[0040] Figure 8 The results of uniaxial tensile tests of PCL and PCL diurea electrospun scaffolds along (Figure 1) and perpendicular to (Figure 2) the main fiber direction are shown.

[0041] Figure 9A The results of uniaxial fatigue tests (10%, 2Hz) on PCL and PCL diurea are shown. Figure 9B The results of pipe fatigue tests using PCL bisurea and PCL UPy implants are shown.

[0042] Figure 10 The results of a curling test on electrospun heart valve implants are shown.

[0043] Figure 11 This shows different curling states of a 3D electrospun heart valve.

[0044] Figure 12 The stent is shown after different stages of the curling of a 3D electrospun heart valve.

[0045] Figure 13 Two images showing DAPI staining of PCL-diurea implants in a sheep model are displayed.

[0046] Figure 14 Two SEM images of PCL and PCL-diurea matrix are shown.

[0047] In the description of this invention and the appended claims, the following terms are used and are interpreted as follows.

[0048] Unless otherwise stated, “polymer” is intended to also include homopolymers, copolymers or supramolecular polymers.

[0049] "Supramolecular polymers" are polymeric arrays of monomeric units aggregated together through reversible and highly oriented secondary interactions, exhibiting polymeric properties in diluted and concentrated solutions as well as in the bulk. The monomeric units of supramolecular polymers themselves do not possess chemically repetitive segments. The orientation and strength of supramolecular bonds are important characteristics of these systems, which can be considered polymers and behave according to existing theories of polymer physics.

[0050] In this application, "supramolecular monomer compound" refers to a compound that forms a supramolecular polymer due to reversible and highly oriented secondary interactions (with other supramolecular monomer compounds).

[0051] Supramolecular polymers are therefore composed of monomers designed to self-assemble into the desired polymeric structure. This contrasts with conventional polymerization reactions, where monomers are linked by covalent bonds. As a result of self-assembly, materials with significantly higher virtual molecular mass are obtained. Examples of supramolecular polymers have been described, for example, in Science, 1997, 278, 1601.

[0052] "Contrast agents" are substances used in medical imaging to enhance the contrast of structures or fluids within the body.

[0053] A "scaffold" is a temporary structure used to support a material (e.g., tissue) during the formation and / or recycling of the material.

[0054] "Structural components" are the parts of a scaffold intended to provide structural properties.

[0055] The “imaging component” is the part of the stent intended to provide imaging properties.

[0056] "Bioactive components" are the parts of a scaffold intended to provide bioactivity.

[0057] The "matrix" is the material on which cells grow.

[0058] Implants are medical devices that replace, support, cover, or enhance existing biological structures that are dysfunctional or damaged.

[0059] The backbone of a polymer is the backbone chain (also called the main chain), which is a series of covalently bonded atoms that work together to create a continuous chain of polymer.

[0060] "Porosity" is measured, for example, by mercury porosimetry, fluid intrusion, and gravimetric analysis. "Pore diameter" is the average size of the openings (pores) in a matrix material. The porosity mentioned in the specification is measured as follows:

[0061] The weight of the support was measured using a balance. Dimensions were also measured (length and thickness for tubes, length, width, and thickness for sheets). Porosity was calculated using the following formula:

[0062] Porosity = (1 - scaffold density / polymer density) x 100%

[0063] The polymer density varies depending on the polymer used, and the stent density is calculated as stent weight / stent volume.

[0064] The term "pore" refers to the space between fibers in a matrix material (i.e., pore size). Pore size and porosity are properties of the scaffold that affect cell connectivity, proliferation, migration, and / or differentiation.

[0065] Minimally invasive surgery is a type of surgery (surgical procedure or other methods) that is less invasive than open surgery.

[0066] Surprisingly, the inventors have discovered that the above-mentioned objectives can be achieved using an implant comprising a matrix material with a porosity of at least 60%. The implant can be surgically inserted into a subject (also known as the recipient of the implant). Because the porosity is at least 60%, the implant can be compressed to minimize its size (also known as curling). Due to the reduced size of the implant, a smaller opening is required for insertion, resulting in less discomfort for the recipient and minimizing recovery time. A porosity of 70% to 90% is preferred. Furthermore, fibrous structures with such porosity allow nutrients to diffuse into the matrix and enable cells to grow inward and / or infiltrate into the matrix.

[0067] The inventors have discovered that they can overcome one of the major limitations currently associated with minimally invasive implantation, such as artificial heart valves, by making them smaller than existing transcatheter heart valves.

[0068] In embodiments of the invention, the implant is a cardiovascular implant, and preferably the implant is selected from (blood) vessels, heart valves, cardiovascular patches, or valvular catheters. Advantageously for the recipient, the implant according to the invention can be implanted via minimally invasive surgery. For these implants, only a small incision is required to facilitate implantation. It is preferred that the implant be applied to the subject (i.e., the patient) via a small incision. Due to the high porosity of the implant, the diameter of the implant can be reduced by at least five times, from a fully unfolded configuration to a fully coiled configuration and from a fully coiled configuration to a fully unfolded configuration.

[0069] In embodiments of the invention, the implant is reinforced by at least one support structure, preferably selected from reinforcing rings, suture rings, or scaffold structures, and preferably biodegradable. The implant is preferably composed of a reinforced matrix material. The presence of the support structure may, for example, be intended to enhance the implant, allow it to curl during minimally invasive surgery, fix it in a precise anatomical position, or allow for repeated needle punctures. Suitable support structures are those well known in the art and are used, for example, for artificial heart valves or as coronary artery stents or stents for other types of arteries. Reinforcement is described, for example, in US 4,626,255, US 6,338,740, US2004 / 0148018, US 3,570,014, and US 4,084,268.

[0070] In embodiments of the invention, the implant has a matrix material composed of a fibrous network. The fibrous network is made of fibers. The fibers enable the implant to have good structural integrity while maintaining its porosity and pores. Preferably, the fibrous network is electrospun fibers. Electrospinning is a technique using a metal target or mold, having a flat, plate-like, or complex three-dimensional form, depending on the desired preform. Polymer fibers are deposited onto the mold by means of an electromagnetic field. The polymer fibers are produced from a solution of one or more polymers in one or more solvents. This electrospinning technique is known in the art and will not be described in further detail in this specification. Dutch patent NL 1026076 (corresponding to US2008 / 0131965) discloses the preparation of articles by means of electrospinning polymer microfibers. The electrospinning setup used in developing such products is climate-controlled and also allows for control of the rotation area, nozzle speed, collector rotation, and the application of smaller negative voltages (up to -4kV). Humidity, temperature, and other factors mentioned above can be used alone or in combination to modify various characteristics of the electrospun fibers. These include, but are not limited to, fiber morphology, fiber diameter, fiber and pore size distribution, porosity, and scaffold thickness.

[0071] In embodiments of the invention, the fibers of the matrix material are composed of one or more supramolecular polymers. By using these types of polymers, the inventors have discovered that implants can be placed into subjects without prior cell seeding. Through the use of supramolecular compounds, the inventors have observed cells attaching, infiltrating, and growing in vivo on and within the implant, while the implant performs the function of the tissue to be replaced or repaired. Therefore, implants can be directly placed into patients. The advantages of direct implantation of scaffolds include reduced generation time and cost, extended storage periods, and eligibility as medical devices, meaning faster regulatory approval. Thus, one or more significant drawbacks of in vitro tissue construct growth or cell seeding are overcome while maintaining the advantages and promise of regeneration, leaving only new, healthy tissue.

[0072] In embodiments of the invention, the implant is biodegradable. This allows the implant to degrade after implantation. Therefore, the implant is replaced by tissue over time. An advantage is that the implant does not require surgical removal, preventing further discomfort for the recipient.

[0073] The inventors have discovered that if one or more supramolecular compounds have a backbone comprising or composed of polycaprolactone (PCL) or a combination of PCL, caprolactone, polylactic acid and / or lactic acid, very good results are obtained regarding the structural characteristics of the implant and the biodegradability.

[0074] In embodiments of the invention, the implant has a matrix material comprising one or more supramolecular compounds, wherein the supramolecular compounds further comprise one or more groups selected from Upy (ureido-pyrimidinone) and / or biuret. Polycaprolactone polymers containing UPy groups are disclosed, for example, in patent application EP 1687378. A preferred supramolecular compound is PCL-biuret (also known as PCLbu). The inventors have found that using these compounds, favorable results can be obtained regarding in vivo tissue growth within the implant and the structural characterization of the implant. Favorable results are obtained with PCL-biuret.

[0075] In the embodiments, one or more supramolecular compounds comprise at least PCL-UPy, preferably with the matrix material composed of PCLUPy. Particularly good results are obtained with PCL-UPy.

[0076] Methods for preparing PCL-bisurea are described, for example, in Chapter 2 of "Biomaterials by the supermolecular control of namofibers" by E. Wisse, ISBN: 978-90-386-1094-8. The chemical formula of PCL-bisurea having a butyl spacer is shown in Formula I below. Alternatively, a spacer may be omitted or another alkyl spacer, such as a hexyl spacer, may be used. PCL acts as a soft block, while the urea group constitutes a hard block. Hydrogen bonds between the hard blocks result in reversible physical linkage.

[0077]

[0078] Where p and n are integers, and p and n > 1. The values ​​of p and n can be varied to obtain different PCL biuret formulations, which may result in different properties. The p value depends on the starting molecular weight of polycaprolactone diol, while the n value relates to the amount of chain extension in the PCL biuret. Preferably, the n value can be in the range of 4–40.

[0079] Methods for preparing PCL-UPy can be described, for example, in Chapter 6 of "Biomaterials by the supramolecular control of nanofibers" by E. Wisse, ISBN: 978-90-386-1094-8. PCL-UPy polymers can be prepared, for example, to contain urea hydrogen-bonded groups (Upy-U 1) or urethane hydrogen-bonded groups (Upy-U 2).

[0080]

[0081] In embodiments of the invention, the fibrous network comprises nanofibers and / or microfibers, preferably with a diameter of 3 to 20 micrometers, more preferably 5 to 10 micrometers. This diameter provides the implant with excellent mechanical and structural stability while offering sufficient porosity for a microstructure that allows cells to grow inward (Balguid, Strategies to optimize engineered tissue towards native human aortic valves, PhD thesis Eindhoven University of Technology, 2008, ISBN 978-90-386-1185-3).

[0082] Nanofibers are fibers with a diameter of less than 1 micrometer. The diameter of microfibers and nanofibers can be obtained by measuring the diameter under a microscope.

[0083] In embodiments of the invention, the matrix material comprises pores having a diameter ranging from 1 to 300 micrometers, preferably from 5 to 100 micrometers. The advantage of these pore sizes is that they allow cells to be cultured to pass through, thus allowing cells to well infiltrate to the full thickness of the preform, which is necessary to ensure tissue formation within the complete preform. The required pore size depends on the size of the cells to be cultured and can be selected accordingly. Human cells are generally larger than animal cells, therefore, there is a difference in the optimal pore size when using animal or human cells.

[0084] In a preferred embodiment of the invention, the matrix material layer has a thickness of at least 100 μm and a maximum of 3000 μm, preferably 200 to 1000 μm. When the implant is composed of matrix material, the implant has a thickness of at least 100 μm and a maximum of 3000 μm, preferably 200 to 1000 μm. The inventors have discovered that at the defined thickness, implants with good structural properties are obtained, thereby enabling these implants to achieve the desired function when implanted in a subject, while simultaneously leading to tissue growth and obtaining tissue of good quality.

[0085] In embodiments of the invention, the implant has a linear elastic stiffness in the range of 0.1-50 MPa, preferably 0.1-10 MPa. The inventors have found that these ranges provide a combination of strength and flexibility necessary to withstand the hemodynamics of the human cardiovascular system. Stiffness values ​​in this range have also been reported in previously published results for natural and tissue-engineered materials (Non-invasive assessment of leaflet deformation and mechanical properties in heart valve tissue engineering, Kortsmit, ISBN: 978-90-386-2002-2 (2009), Stradins et al. (2004), Clark, (1973)).

[0086] In embodiments of the invention, the implant has at least 30%, preferably at least 45%, and most preferably at least 60% of a (linear) elastic regime (linear elastic stiffness has been measured using a standard uniaxial tensile test (for description, see ISO 13934-1:1999 Textiles—Tensive properties of fibrous fabrics—Part 1: Determination of maximum force and elongation at maximum force using the band method)). This ensures that the implant according to the invention does not exhibit plastic deformation or breakage (strain applied to the implant in vivo) within the physiological strain zone. For example, for natural heart valves, physiological strain of ~60% has been reported (Billiar & Sacks, (2000), Driessen et al., 2005)). Existing prosthetic bioprosthetic valves fail to meet these values, with maximum strains observed in glutaraldehyde-treated porcine valves during diastolic pressure increase at 2–4% and 3–10% in the circumferential and radial directions, respectively (Adamczyk & Vesely, 2002), and average strains in the heart valve leaflets at 4–10% (Sun et al., 2005). Furthermore, chemically fixed anisotropic tissues have been described as becoming more isotropic than fresh tissues due to chemical crosslinking (Zioupos et al., 1994) and exhibiting lower compliance (Broom et al., 1982; Schoen et al., 1997; Billiar & Sacks, 2000). Commercially available polymers exhibiting short elastic phases in uniaxial tensile tests have also shown inadequate performance in in vitro valve testing settings. Therefore, an extended elastic phase is important for implants intended for in vivo use and for achieving the function of replacing or repairing tissue.

[0087] In the implant embodiment of the present invention, the fibers (of the matrix material) have a preferred orientation. Preferably, the fibers in the implant are arranged in such a way that they are aligned in a direction substantially perpendicular to blood flow when the implant is inserted.

[0088] Preferably, the fibers are arranged circumferentially along the imaginary axis of the implant, wherein in the case of a tubular implant, the axis points in the direction of blood flow. In the case of a heart valve leaflet, the fibers are preferably arranged in accordance with... Figure 1 Arranged in the same way as described in the text.

[0089] Such orientations can be introduced during implant fabrication (i.e., electrospinning). These fibrous structures mimic the natural fiber arrangement in natural tissues, such as the hammock-like collagenous architecture in natural heart valves (as in...). Figure 1 As described in [reference needed], Sauren et al. (1981) and the orientation of helical collagen fibers in natural arteries (as described in [reference needed]). Figure 2 As described in [the text], by Holzapfel (2000).

[0090] By simulating the extracellular matrix of the natural environment, tissues can grow to have good structural properties and eventually develop into natural-like structures.

[0091] In embodiments of the invention, the ratio of linear elastic stiffness between stiffness in the preferred fiber direction and stiffness perpendicular to the preferred fiber direction is at least 2:1, preferably at least 4:1, more preferably at least 10:1, and even more preferably at least 50:1. Implants with such ratios have good structural properties while still providing a matrix for cell growth in a simulated natural environment.

[0092] In embodiments of the invention, the implant further comprises a bioactive compound and / or a contrast agent. The contrast agent may be present in the implant to monitor how rapidly degradation occurs within the implant and to determine the ultimate success of the tissue engineering procedure. For example, the contrast agent may be visible in relevant clinical imaging techniques such as computed tomography (CT), magnetic resonance imaging (MRI), and / or ultrasound diagnostics (or ultrasound examination) using ultrasound for imaging purposes. Preferred contrast agents are described in European Patent Application No. 10193654, which describes fluorinated polymers having a glass transition temperature (Tg) below 40°C, preferably below 20°C, more preferably below 0°C, as imaging markers or contrast agents in 19F magnetic resonance imaging (MRI). Based on the total mass of the polymer, the amount of fluorine (19F) in the fluorinated polymer is preferably at least 5 wt%. The fluorinated polymer comprises at least one polymer selected from the group consisting of (per)fluorinated polyethers, (per)fluorinated polyesters, (per)fluorinated poly(meth)acrylates, and (per)fluorinated polysilicone, preferably (per)fluorinated ethers. The polymer may be incorporated into the polymer constituting the fibrous network. Optionally, the polymer exists alone in the fibrous network.

[0093] Bioactive compounds can be added to promote, for example, cell infiltration, retention, differentiation and proliferation, as well as tissue formation and remodeling.

[0094] The present invention also relates to a method for preparing an implant having a matrix material comprising one or more supramolecular compounds, preferably an implant as described above, wherein the matrix material has a porosity of 60%, preferably between 70 and 90%, the method comprising the following steps:

[0095] -Provide molds;

[0096] - Applying matrix materials to a mold using electrospinning with the aid of one or more supramolecular compounds; and

[0097] - Separate the matrix material from the mold.

[0098] The resulting implant has the same advantages as described above. In embodiments of the invention, the method further includes the step of providing at least one support structure to the implant or matrix material.

[0099] The present invention also relates to a coiled implant, preferably an implant according to the invention as previously described, which can be coiled to a diameter size of up to 20% compared to the diameter of the implant before coiling (uncoiled). This allows the implant to be easily delivered via minimally invasive surgery. A further advantage is that, since the implant can be coiled to 20% of its initial diameter, it can be delivered to subjects currently excluded from receiving implants via minimally invasive surgery because, for example, their arteries are too narrow for a coiled implant according to the prior art to pass through. Preferably, the coiled implant is a heart valve implant. The coiling of the implant according to the invention is achieved by methods known in the art. Examples of coiling are given in the embodiments.

[0100] The present invention also relates to a method for growing a valve, comprising the step of providing an implant, preferably an implant according to the invention, to a subject (patient). The implant is preferably an implant according to the invention. This step may be preceded by a step of making an incision in the subject's skin.

[0101] After the implant is placed in a patient (human or animal), the implant is able to function as tissue to grow both after implantation and before inward cell growth occurs, wherein inward cell growth occurs after implantation and the matrix material degrades over time.

[0102] The invention will be further illustrated by the following non-limiting examples. The appended claims also form part of the description of this application. Example

[0103] Example 1

[0104] The implant is prepared according to the following method. The required amount of PCL, PCL diurea, or PCL UPy is dissolved in a suitable solvent / solvent mixture and stirred until dissolved. The resulting solution is delivered at a constant flow rate (flow varying over time is also possible and will produce scaffolds with different properties) to an electrically charged nozzle, typically using an injection pump. A high voltage is applied to the nozzle. The voltage difference used (a combination of positive and negative voltages) is 10-20 kV, although other voltages are possible for fiber generation. A rotating collector, typically in cylindrical form, is placed. The collector is connected to the ground or a negative terminal. The rotation speed is typically 100 rpm. Fibers are deposited onto the collector. The length and thickness of the resulting implant are affected by flow, voltage, collector rotation speed, scanning, and nozzle speed. Once the desired thickness is reached, rotation is stopped and the collector is removed. The implant in the collector is vacuum dried and annealed (-37°C) overnight. The implant is removed via the collector by immersing it in warm water (-37°C), although other removal methods are also possible. SEM images of electrospun fiber mesh are shown in Figure 14 middle.

[0105] Example 2

[0106] Implants obtained according to the method of Example 1 were tested according to ISO 5840:2005. The hemodynamic performance of the valve implants was evaluated by loading physiologically relevant flow and pressure using a simulated circulatory system. Figure 7Images of the simulated circulation used are shown. The valves installed in this simulated circulation are subjected to either pulmonary artery pressure and flow or aortic pressure and flow. The fluid used in the experiment was a saline solution. In the simulated circulation system shown above, the circulating fluid is replaced by a computer-controlled piston pump (P). The piston is connected to a servo motor (SM) system, which is controlled by a motion control panel. The piston fills the left ventricular cavity (LV; VLV = 1L) from the reservoir through the model mitral valve (MV), and then injects fluid through the arterial valve (AV) into the WindKessel (WK) model, which consists of two resistors (R1 and R2) and a compliance tank (C; VC = 2L). The fluid flows back to the reservoir from the WindKessel model through a portion of the silicone tubing. The pulsating flow through the arterial valve is measured by a flow meter (FM1). Furthermore, the mean cardiac output is recorded using a clamp-on flow sensor (FM2) at the silicone tubing outlet of the WindKessel. Ventricular and arterial pressures (Pv and Pa) were recorded using pressure sensors connected to a bridge amplifier (Picas, Peekel Instruments). Signals were recorded via a data acquisition board and stored on the PC's hard drive. An endoscope was introduced into the systemic artery, and a high-speed color camera (M5) was connected to the endoscope to capture valvular dynamics at a frame rate of 200 Hz throughout the entire cardiac cycle. Valve pressure and flow were monitored over time to assess valvular functionality.

[0107] Valves prepared from PCLbu and PCL were subjected to systemic conditions at 120 / 80 mmHg for 20 hours. The results clearly show that PCLbu ( Figure 3 and 4 Implants have advantages over PCL implants ( Figure 5 and 6 Better results.

[0108] Figure 4 and 6 The images show the valves in open (top left) and closed (top right) configurations at the start of the test, and in open (bottom left) and closed (bottom right) configurations after 20 hours. The images of the tested valves clearly show that the PCL valve was damaged, while the PCLbu valve remained largely undamaged.

[0109] Example 3 Uniaxial Tensile Test

[0110] The Zwick tensile stage is connected to a measurement module, which is connected to a PC. After necessary preparation, the system can be used to retrieve stress-strain characteristics, for example, of biological tissue strips. First, the sample is held in place by two clamps during the tensile stage. Software on the PC triggers the tensile stage, thus stretching the sample. During this stretching, displacement and force are recorded and subsequently stored in a TRA file. Based on this file, which includes sample dimensions, mechanical parameters (Young's modulus, ultimate tensile strength / strain) can be determined using Matlab.RT (dry), grip-to-grip separation = 9 mm, elongation = 9 mm / min, and a 20 N load cell.

[0111] The Young's modulus [Pa] is calculated using the width and thickness, and it is a measure of the strip stiffness.

[0112]

[0113] F[N] represents "force" and l[m] represents the length of the sample (length[mm]*10). -3 ).

[0114] AO[m 2 Indicates the cross-sectional area of ​​the sample before testing (width [mm] * thickness [mm] * 10). -6 ).

[0115] Figure 8 The results of uniaxial tensile tests on PCL and PCL-bisurea electrospun scaffolds along (left) and perpendicular (right) the major fiber direction are shown, revealing a significant difference in the elastic phase (linear) between the two materials. For PCL, plastic deformation and breakage were observed at approximately 10% strain and below 15% strain (perpendicular). For PCL-bisurea, no plastic deformation or breakage was observed in the physiological strain zone (<60%) under the same conditions.

[0116] Example 4 Fatigue Test

[0117] A) Uniaxial fatigue test

[0118] Using a setup similar to that used in uniaxial tensile testing, cyclic loading was performed on PCL and PCL-diurea to evaluate fatigue behavior in uniaxial fatigue testing (10% strain, 2 Hz). The results confirmed the timeline and failure modes observed in the valve test (Example 2), showing fatigue within 1000 cycles (in vivo <1 / 2 hour) for valves made from PCL, while no fatigue was observed for valves made from PCL-diurea. Even after 1 million cycles (in vivo >11 days), no fatigue damage was observed for the PCL-diurea valve. Figure 9A ).

[0119] B) Catheter fatigue test

[0120] The fatigue resistance of implants obtained according to the method of Example 1 was tested by applying cyclic pressure loads to tubular stents. This benchtop test was used to determine the durability of vascular devices by subjecting them to hydrodynamic pulsating loads with a relevant pressure differential. Device samples were immersed in an environmental chamber for 3 million fatigue cycles. Test conditions were designed to meet the requirements of in vitro mechanical fatigue testing as described in ASTM F 2477-07 "Standard Test Methods for in vitro Pulsatile Durability Testing of Vascular Devices" and FDA Guidance 1545 (2010) "Non-Clinical Engineering Tests and Recommended Labeling for Intravascular Devices and Associated Delivery Systems". The test was controlled and monitored via a pressure control method, which specifies that the cyclic pressure range be controlled within the desired range. The frequency was set to 5 Hz, the temperature to 37°C, and the pressure range to 35 / 15 mmHg. As an output, the outer diameter of the stent under test was measured as a function of the number of cycles.

[0121] Figure 9B It is clearly shown that, compared to UPy material, the outer diameter of the PCLbu-based material increases more significantly as a function of cycle number, indicating that UPy material has greater resistance to fatigue failure. Figure 9A The results shown are compared in Figure 9B As can be clearly seen below, PCL-Dureur exhibits better fatigue resistance than PCL implants. Furthermore, PCL-UPy implants exhibit even better fatigue resistance than PCL-Dureur implants.

[0122] Figure 9B Results for two different PCLbu configurations and two different UPy configurations are also shown, demonstrating that the fatigue properties of PCLbu-based and UPy-based scaffolds can be improved by adjusting the scaffold material.

[0123] Example 5

[0124] The Young's modulus of isotropic and anisotropic electrospun PCLbu strips was obtained, and the stiffness ratio was calculated from them. The results are listed in Table 1.

[0125] Table 1

[0126]

[0127] Table 1 clearly shows that anisotropically rotated PCLbu strips exhibit both high Young's modulus (when measured parallel to the preferred fiber orientation) and low Young's modulus (when measured perpendicular to the preferred fiber orientation). Table 1 also shows that for isotropically rotated PCLbu strips, the Young's modulus measured in the parallel direction is comparable to that measured in the perpendicular direction.

[0128] Example 6: Curling Test

[0129] A curling test was performed on an implant according to the invention with a diameter of 28 mm. The implant was shrunk to a certain diameter using a commercially available device for curling transcatheter heart valves. The different stages of the curling method are described in... Figure 11 In the middle. Then return them to their normal size. Measure the diameter again after unfolding. The results are shown in Figure 10 In the diagram, the striped bars represent the diameter of the implant when it is curled up, the white bars represent the diameter of the implant after it has been returned to its uncurled size, and the black bars represent the same diameter as the white bars but after the implant has been immersed in PBS at 37°C for 1 hour.

[0130] The results clearly show that even after the implant had been curled to a diameter of 6 mm, it was observed to return to its initial size (29 mm in diameter) after being immersed in PBS at 37°C for 1 hour. Figure 12 The image shows the valves after they have returned to their uncurled size. The numbers next to the image correspond to... Figure 10 The number.

[0131] Example 7 In vivo testing

[0132] The implant according to the invention was implanted in the pulmonary valve location of an adult sheep using open-heart surgery. Cells were visualized by DAPI staining, and cell infiltration was assessed under a microscope. Cell infiltration was assessed at 1 day, 7 days, and 8 weeks post-implantation. Results are described in... Figure 13 The top left image shows cell infiltration in the PCL-Upy implant after 1 day, the top right image shows cell infiltration after 7 days, and the bottom image shows cell infiltration in the PCL-Upy implant after 8 weeks. The image clearly shows intracellular infiltration in both the PCL-Upy and PCL-Upy implants.

Claims

1. A coiled implant comprising a matrix material, said matrix material being composed of a fibrous network obtained by electrospun fibers, and said fibrous network being composed of one or more supramolecular compounds, said fibers being generated from a solution of one or more polymers in one or more solvents, said fibrous network having a porosity of at least 60% and comprising pores in the diameter range of 1-300 micrometers, and said fibrous network having a linear elastic stiffness of 0.1-50 MPa.

2. The implant of claim 1, wherein the implant is used to generate tissue in vivo.

3. The implant of claim 1, wherein the implant does not contain cells prior to implantation.

4. The implant of claim 1, wherein the implant can be rolled up to a diameter of up to 20% compared to the unrolled implant.

5. The implant according to claim 1, wherein the implant is a cardiovascular implant.

6. The implant of claim 5, wherein the implant is selected from blood vessels, heart valves, cardiovascular patches, or valvular catheters.

7. The implant of claim 6, wherein the vascular vessel is a blood vessel.

8. The implant according to any one of claims 1-7, wherein the implant is reinforced by at least one support structure.

9. The implant according to claim 8, wherein the at least one support structure is selected from a reinforcing ring, a suture ring, or a scaffold structure.

10. The implant of claim 8, wherein the at least one support structure is biodegradable.

11. The implant of claim 8, wherein the implant is composed of a reinforcing matrix material.

12. The implant according to any one of claims 1-7, wherein the fibrous network is biodegradable.

13. The implant of claim 1, wherein the one or more supramolecular compounds have a backbone comprising or composed of: polycaprolactone (PCL); or a combination of PCL, caprolactone, polylactic acid and / or lactic acid.

14. The implant of claim 13, wherein the supramolecular compound further comprises one or more groups selected from UPy (ureidopyrimidinone) and / or diurea.

15. The implant of claim 14, wherein the one or more supramolecular compounds comprise at least PCLUPy.

16. The implant of claim 14, wherein the fibrous network is composed of PCL UPy.

17. The implant according to any one of claims 1-7 or 13-14, wherein the fibrous network comprises nanofibers and / or microfibers.

18. The implant of claim 17, wherein the diameter of the fiber is 3 to 20 micrometers.

19. The implant of claim 18, wherein the diameter of the fiber is 5 to 10 micrometers.

20. The implant according to any one of claims 1-7 or 13-14, wherein the fibrous network comprises pores with a diameter in the range of 5-100 micrometers.

21. The implant according to any one of claims 1-7 or 13-14, wherein the fibrous network has a linear elastic stiffness in the range of 0.1-10.0 MPa.

22. The implant according to any one of claims 1-7 or 13-14, having a linear elastic phase of at least 30%.

23. The implant according to any one of claims 1-7 or 13-14, having a linear elastic phase of at least 45%.

24. The implant according to any one of claims 1-7 or 13-14, having a linear elastic phase of at least 60%.

25. The implant of claim 12, wherein the fibers have a preferred orientation.

26. The implant of claim 25, wherein the preferred orientation is circumferential.

27. The implant of claim 25, wherein the ratio of linear elastic stiffness between the fibers in the preferred orientation direction and the fibers perpendicular to the preferred orientation direction is at least 2:

1.

28. The implant of claim 25, wherein the ratio of linear elastic stiffness between the preferred orientation direction and the direction perpendicular to the preferred orientation direction is at least 4:

1.

29. The implant of claim 25, wherein the ratio of the linear elastic stiffness between the preferred orientation direction and the direction perpendicular to the preferred orientation direction is at least 10:

1.

30. The implant of claim 25, wherein the ratio of the linear elastic stiffness between the preferred orientation direction and the direction perpendicular to the preferred orientation direction is at least 50:

1.

31. The implant according to any one of claims 1-7 or 13-14, wherein the implant further comprises a bioactive compound and / or a contrast agent.

32. The implant according to any one of claims 1-7 or 13-14, wherein the implant is used to promote the growth of cardiovascular tissue in a patient, and wherein the implant is cell-free prior to implantation.

33. The implant according to any one of claims 1-7 or 13-14, wherein the fibrous network is free of bioactive compounds.

34. A method for preparing a coiled implant having a matrix material composed of a fibrous network, the fibrous network being obtained from electrospun fibers, and the fibrous network being composed of one or more supramolecular compounds, and the fibers being produced from a solution of one or more polymers in one or more solvents, wherein the fibrous network has a porosity of at least 60% and contains pores with diameters in the range of 1-300 micrometers, and the fibrous network has a linear elastic stiffness of 0.1-50 MPa, the method comprising the following steps: - Mold provided: - The fibrous network is applied to a mold by electrospinning with one or more supramolecular compounds to obtain an implant; - Separate the implant and the mold; and - Curl the implant.

35. The method of claim 34, further comprising the step of providing at least one support structure to the implant.

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