Method for preparing a three-dimensional medical scaffold

Abstract

1. A method for preparing a sterile scaffold for medical use, comprising the steps of: The method includes the steps of: i) filling a fiber mesh containing fibers of a polylactide polymer or copolymer (commonly referred to as PLA) with collagen to obtain a PLA-collagen scaffold; ii) drying the PLA-collagen scaffold obtained from step i); and iii) sterilizing the PLA-collagen scaffold obtained from the dried step ii) to obtain a sterile scaffold, which has improved biomechanical properties compared to a non-sterile scaffold.

Description

[Technical field]

[0001] The present invention relates to a method for preparing a three-dimensional scaffold for use as a pharmaceutical or cosmetic biomaterial. The scaffold is composed of polylactide or polylactic acid polymer or copolymer (commonly designated PLA) and collagen. The method includes steps that allow the preparation of a scaffold with the desired properties in terms of i) a set of biomechanical characteristics, ii) stability, and iii) purity. Furthermore, the method does not include steps that may result in significant degradation of the components used, namely PLA and collagen, or in significant degradation of the resulting PLA-collagen scaffold. Moreover, it has been unexpectedly shown that a sterilization step confers certain biomechanical properties to the scaffold. [Background technology]

[0002] The need to develop tissue substitutes and regeneration platforms is one of the most demanding and challenging applications in modern tissue engineering. Three-dimensional biomaterial structures (scaffolds) are highly desirable, matching the biomechanical properties of tissues and closely mimicking in vivo behavior (facilitating cell adhesion, growth and tissue formation). Such biomaterials usually help the body to reconstruct damaged tissues, ultimately minimizing the associated pain and healing time. Especially for tissues where load-adaptive properties are required, the combination of static and dynamic biomechanical properties of the scaffold is crucial for the ultimate success of the treatment. Every advance in the development of scaffolds should ensure a high correlation between in vitro conditions and the expected in vivo tissue regeneration. The non-toxic biodegradation of the scaffold should gradually transmit stresses to the new growing tissue over an appropriate period of time. The correct mechanical stimulus synergy is highly dependent on the scaffold material, its intended biological environment and the presence of cells.

[0003] One of the most challenging applications of biomedical scaffolds is articular cartilage (AC) repair. Damage and degeneration of AC is ongoing even in young and active populations, not only due to age, obesity or systemic diseases, but also due to physical causes such as injuries. If untreated, these defects can progress to osteoarthritis (OA), which affects more than 240 million people worldwide. Natural wound healing of full-thickness cartilage defects often results in the formation of fibrocartilage, which is functionally and biomechanically inferior to the original hyaline cartilage, further deteriorating the tissue and predisposing the joint to osteoarthritic changes. A vicious cycle that has been initiated will eventually require total or partial joint replacement. Therefore, a biomaterial solution with the ability to regenerate cartilage is highly desirable to treat cartilage lesions at an early stage before the manifestation of OA.

[0004] Clinically used biomaterials include a variety of naturally derived and synthetic materials. The advantage of natural materials is their natural feasibility for the purpose, while the application of animal-derived materials (xenografts) includes certain risks such as contamination and undesired immune responses. This can be avoided by using synthetic materials that do not cause foreign body or hypersensitivity reactions. Synthetic materials can be made biologically more favorable and biocompatible. On the other hand, compared to naturally derived materials, synthetic polymers usually lack the desirable intrinsic biological cues that promote cell adhesion, proliferation and tissue recovery. However, any biomaterial is always difficult to evaluate and optimize for clinical use and for the purpose of aiming at "precision medicine" solutions. Currently, it is widely expected that the current level for evaluating the mechanical functions of biomaterials and tissue engineered constructs is still insufficient.

[0005] Synthetic materials of fibrous origin are often used for AC repair applications. These scaffolds have a porosity of 75–85%.

[0006] The structure, function and biomechanical behavior of AC are highly complex, anisotropic and time and loading history dependent. Articular cartilage consists of a relatively small number of chondrocytes surrounded by a multicomponent matrix, which can be imaged as a complex of 70–85% water and the remainder proteoglycans (proteins formed by glycosaminoglycans linked together in bottle-brush-like structures) and collagen. Proteoglycan and water concentrations vary with depth in the cartilage tissue.

[0007] Without blood supply and lymphatic drainage, articular cartilage remains isolated and virtually devoid of the wound healing response of other connective tissues. High exposure of the tissue to biomechanical abnormalities results in high incidence levels of cartilage lesions. Such lesions often develop into osteoarthritis (OA), due to traumatic or prolonged non-physiological loading. OA is the largest cause of musculoskeletal disease in the world, with an incidence rate of 7–10% in the Western population. The estimated cost of OA in newly diagnosed patients is $6,800 per year, so postponing OA for 10 years would result in cost savings of $68,000 per patient. Expenditures for OA in the EU are approximately 15–20 billion per year. Although not traditionally indicated for the treatment of OA, cartilage repair has attracted growing interest due to its potential to alter the progression of the degenerative disease in the hope of delaying or avoiding the need for joint replacement surgery.

[0008] In addition to the significant morbidity and potential for disability, cartilage trauma and degeneration also have a large economic impact. Estimating the annual incidence of cartilage lesions to be 23 per 100,000 population, there are more than 100,000 patients in the EU with cartilage defects in the knee requiring restorative treatment. The prevalence of cartilage pathologies is expected to increase rapidly in the coming decades due to an ageing population and increasing rates of obesity; the demand for knee replacement surgery is predicted to increase significantly by 2030.

[0009] On the other hand, young patients with symptomatic cartilage lesions represent a challenging population due to a combination of high functional demands and limited treatment options. The goal of articular cartilage repair treatment is to restore and maintain normal function of the joint with a repair tissue structure indistinguishable from natural hyaline cartilage. However, current repair techniques for cartilage lesions are insufficient and require development.

[0010] After surgical repair, the biomechanical properties of the repair site are weakened, necessitating a reduction in postoperative loading. Thus, the lack of mechanical stimulation is expected to delay tissue turnover and healing; therefore, recovery times remain long. As mentioned above, biomaterial scaffolds can provide structural support to the healing lesion to allow early load-bearing and thus enhance the healing process. A wide variety of three-dimensional scaffolds, both natural and synthetic, have been introduced for cartilage repair.

[0011] One type of scaffold of particular interest in the context of the present invention is a scaffold made from a polylactide polymer or copolymer, commonly designated PLA. Such scaffolds have been the subject of several publications:

[0012] PLA scaffolds are described, for example, in Non-Patent Document 1. Gamma-irradiated PLA scaffolds were immersed in a solution of recombinant human collagen type II to obtain rhCo-PLA scaffolds, which were tested in a porcine study using membrane-induced autologous chondrocyte implantation (MACI). Results were compared with a MACI treatment group using a commercially available membrane (Chondro-Gide®) and an untreated control group. Both treatment groups showed improvement compared to the control group, but the differences were not statistically significant. In several rhCo-PLA-treated lesions, the mechanical properties, as well as the repair tissue structure, were very similar to those of healthy cartilage.

[0013] Non-Patent Document 2 relates to a sterilized PLA scaffold doped with a solution of recombinant human collagen III to obtain a rhCo-PLA scaffold. Two types of scaffolds were compared: PLA and rhCo-PLA scaffolds: the results of the biomechanical comparison of the two scaffold types showed how both collagen addition and medium composition altered the elastic, viscoelastic and inelastic properties of the scaffold. In this study, the analysis of the in vivo cartilage repair scores from the study by Muhonen et al. was additionally carried out using BUGS-Bayesian inference Using Gibbs Sampling, a form of Markov Chain Monte Carlo sampling. The results using a normal or Poisson distribution of the sum of normalized ICRS (International Cartilage Repair Society) scores showed that rhCo-PLA had a statistically significantly higher mean score (0.515) compared to 0.38 for the commercial scaffold and 0.288 for the natural healing control group.

[0014] Non-Patent Document 3 relates to cell therapy in combination with biomaterial scaffolds. The experimental scaffold type used was based on recombinant human collagen-polylactide (rhCo-PLA) prepared by immersing a sterile PLA scaffold in a solution of recombinant human collagen II or III. The control scaffold type was a commercially available Chondro-Gide® membrane made from type I / III collagen of porcine origin. The results showed that type II collagen promotes the proliferation of mesenchymal stromal cells (MSCs), but the collagen type used in the rhCo-PLA scaffold does not affect MSC differentiation during in vitro culture, and that chondrogenic differentiation of MSCs leads to cell hypertrophy in the rhCo-PLA scaffold and the commercially available collagen membrane. However, it was said that the limitations of static cell culture may have influenced the results, since hypertrophy is a common phenomenon in in vitro chondrogenesis of MSCs, whereas MSCs in in vivo articular cartilage defects do not show upregulation of hypertrophy markers. In the joint, cells are subjected to cyclic mechanical loading that provides nutrient flow and differentiation cues to the joint.

[0015] US Patent No. 5,399,993 (University of Helsinki et al.) relates to a three-dimensional material obtained by immersing a sterile felt of a polymer in a solution of collagen. The polymer may be PLA. Such scaffolds have been shown to have improved retention / stiffness compared to scaffolds containing bovine collagen type I or recombinant human collagen type II. rhCo-PLA showed better performance after implantation in the right knee of a pig after 4 months of healing compared to Chondro-Gide® scaffold, a two-layered hydrophilic collagen type I / III membrane extracted from a pig.

[0016] However, there is a clear need for a cost-effective, safe and reliable scaffold that has the desired properties yet is safe and sterile. Moreover, it would be advantageous if such a scaffold had general utility not only in cartilage repair but also in medicine or cosmetics in general, including human and veterinary use. [Prior art documents] [Patent documents]

[0017] [Patent Document 1] WO2016 / 042211 [Non-patent literature]

[0018] [Non-Patent Document 1] Muhonen et al., Wiley Online Library DOI 10.1002 / jor 23099, published online in 2015 [Non-Patent Document 2] Gasik et al., Bioengineering and Biotechnology, Volume 6, November 2018 [Non-Patent Document 3] Salonius et al., J.Cell.Physiol.2019 Summary of the Invention [Means for solving the problem]

[0019] The present invention relates to a method for preparing a sterile scaffold for medical use, comprising the steps of: i) filling a fiber mesh containing fibers of polylactide or copolymer (commonly referred to as PLA) with collagen to obtain a PLA-collagen scaffold; ii) drying the PLA collagen scaffold obtained from step i); iii) sterilizing the PLA-collagen scaffold obtained from drying step ii) to obtain a sterile scaffold.

[0020] During preparation, measures are taken to avoid unnecessary degradation of the PLA, collagen and PLA-collagen scaffolds. Thus, the step of sterilizing the PLA-collagen scaffolds is carried out without subjecting the scaffolds to any temperature increase, or at least only a small temperature increase. This can be done, for example, by choosing a sterilization method and / or by taking special precautions to avoid temperature increase. Furthermore, as shown in Example 3 herein, sterilized scaffolds have improved biomechanical properties compared to non-sterile scaffolds, which in turn also improves the biomechanical properties compared to scaffolds prepared under aseptic conditions.

[0021] The fiber mesh used in the method of the present invention comprises: i) providing PLA in solid form; ii) subjecting the PLA to a process, thereby obtaining fibers of PLA; and iii) subjecting the obtained fibers to a process, thereby obtaining a mesh of fibers.

[0022] PLA fibers can be produced by known methods. In an embodiment of the present invention, a preferred method for producing PLA fibers is by spinning, such as melt spinning or electrospinning. Spinning is typically performed by melting PLA and subjecting the molten PLA to a spinning process, or dissolving PLA in a suitable solvent and subjecting the PLA solution to a spinning process. The resulting fibers may be subjected to a process whereby the fibers are included in a mesh, and such a mesh may be subjected to a process whereby a 3D structure is obtained; the mesh may be subjected to a process that ensures that the fibers of the mesh or its 3D structure are preserved.

[0023] Scaffolds of different structures can be obtained. 1. One-dimensional (1D) fiber structure where collagen is packed onto the fibers (length and diameter of the fibers can vary) 2. Impermeable two-dimensional (2D) substrates: these structures allow the inclusion of bioactive entities (e.g. cells) in the structure. Typically, cells are cultured in a 2D environment. 3. Three-dimensional (3D), nanoporous hydrogel scaffolds, collagen is usually located on top of the hydrogel (i.e., interacting with a 2D substrate with a nanoscale surface) or encapsulated inside the 3D structure (cells must degrade the surrounding hydrogel to migrate or extend the process). 4. Three-dimensional (microporous) scaffolds, which have a high degree of porosity (typically >70%), allowing bioactive entities to extend in three dimensions; depending on the pore size of the scaffold, three-dimensional scaffolds can align along one-dimensional scaffold struts or attach to multiple struts and extend in three dimensions.

[0024] In the context of the present invention, the description of a scaffold is also intended to encompass sterile scaffolds.

[0025] Generally, the scaffold of the present invention has a three-dimensional structure. As mentioned above, the three-dimensional structure may further be in the form of a nanoporous hydrogel scaffold, where the bioactive entity is on top of the hydrogel or is encapsulated inside the 3D structure and the surrounding hydrogel must be degraded to release the bioactive entity, or the three-dimensional structure may be in the form of a microporous scaffold, where the bioactive entity can extend in three dimensions due to the high porosity of the scaffold.

[0026] The scaffold of the present invention may be used as is or may be loaded with one or more bioactive entities. The bioactive entities may be one or more cells or may be drugs that are active in a biological environment, particularly in the mammalian body. The cells loaded into the scaffold of the present invention may be cells intended to repair diseased or damaged tissue. Other bioactive entities may be medicinal substances suitable for relieving pain or suitable for treating diseases in a particular tissue. They may be medicinal substances intended for systemic use, but it is easy to administer medicinal substances to the implant. The scaffold may be in the form of an implant or a dressing. The scaffold may be used in medicine, such as human medicine and veterinary medicine. The scaffold of the present invention may be used for different medical purposes, particularly in connection with the repair of cartilage, such as AC, or in connection with osteochondral repair. The scaffold of the present invention may also be used in treatment regimes, such as AO.

[0027] While natural cartilage is not porous, for example, the porosity of natural bone, which contains a large amount of collagen, ranges from 50 to 90% depending on the type of bone. Tissue engineering scaffolds require a fairly porous, interconnected, open pore structure to ensure tissue ingrowth and flow transport of nutrients and metabolic waste products. The porosity of a material can be determined by multiple means. For example, micro-computed tomography (micro-CT) analysis, image analysis (such as scanning electron microscopy or transmission electron microscopy), gas pycnometry as well as mercury and liquid extrusion porosimetry can be used. For overall porosity, micro-CT analysis is considered the best method since it provides reliable results regarding the determination of overall porosity.

[0028] The fiber mesh used in the method of the present invention has a porous structure. The porosity of the PLA fiber mesh is about 80 to 99%, preferably about 85 to 95%. The porosity of the scaffold having both the PLA fiber mesh and the collagen component is about 70 to 99%, preferably about 80 to 95%. The scaffold before and after sterilization has the same range of porosity. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0029] As mentioned above, the present invention provides a method for preparing a sterile scaffold, the method comprising: i) filling a fiber mesh containing fibers of polylactide or copolymer (commonly referred to as PLA) with collagen to obtain a PLA-collagen scaffold; ii) drying the PLA collagen scaffold obtained from step i); iii) drying and sterilizing the PLA collagen scaffold obtained from step ii) to obtain a scaffold.

[0030] The method of the present invention is designed to avoid any unnecessary degradation of the PLA and collagen used and to avoid unnecessary degradation of the fiber mesh and / or PLA-collagen scaffold. Thus, the amount / number of monomers present in the starting material PLA does not change significantly during the method, so that the amount / number of PLA monomers in the final scaffold is close to the starting PLA material. During the pre-process steps of the present invention, PLA is subjected to high temperatures to obtain a fiber mesh containing PLA fibers, and some degradation of the raw PLA material is expected, for example, during the spinning process to obtain a mesh of PLA fibers. Generally, the monomer content of the raw PLA material (before fiber formation) is very low, such as about 0.1%, and the monomer content of the fiber mesh is usually about 1% at most.

[0031] Similarly, only minimal degradation of collagen occurs during the manufacture of the scaffold by the method of the present invention. The stability of the scaffold during its transformation into a sterile scaffold is ensured, inter alia, by avoiding temperature rise of the collagen component during manufacture, especially during the sterilization process, and by ensuring that the sterilization method does not have a widespread negative effect on the biomechanical properties of the scaffold. As previously mentioned herein, the sterilization process has an unexpectedly positive effect on the biomechanical properties of the scaffold. Thus, sterilization by gamma radiation at room temperature or low temperature confers a more stable structure in both the dry and wet states of the scaffold. The scaffold becomes stiffer and has less variation in biomechanical properties, which is advantageous in terms of the approval of the prepared scaffold.

[0032] Regarding stability issues, PLA contains ester bonds that are susceptible to heat, radiation and hydrolysis. Collagen can also be degraded by several mechanisms, including enzymatic, radiation-induced or temperature-dependent degradation.

[0033] The resulting scaffold may be used as is, or may be loaded with bioactive entities such as, but not limited to, cells, or may be loaded with drug substances and used as an implant. In such cases, sterilization of the scaffold may be performed before or after loading the scaffold with bioactive entities such as, but not limited to, drug substances or cells.

[0034] The methods of the present invention provide scaffolds having one or more of the following properties: When used for implants, the scaffold should be biocompatible to support cell viability and biodegradable so that new tissue will replace the scaffold over time and degradation products will not be harmful locally or systemically. The scaffold should have a highly porous 3D structure with interconnected pores so that cells can migrate within the scaffold and so that cells can grow and form new tissue. · It should have adequate porosity to load cells into the scaffold before use so that the cells can remain within the scaffold and proliferate to form new tissue. For example, to ensure that mechanical stimulation works well around the AC and allows new tissue formation, the scaffold should have a biomechanical strength appropriate for its application. While a scaffold cannot closely mimic the native tissue it is intended to replace, it should have the compressive and decompressive capabilities appropriate for the application, as well as fluid absorption and interstitial fluid pressurization during filling and compression. · The scaffold should have an appropriate degradation time such that the application can support tissue ingrowth and subsequent formation of new tissue.

[0035] Polymers used in the method of the present invention A PLA-collagen scaffold is prepared according to the present invention. In the context of the present invention, PLA is intended to include all stereoisomeric forms of polylactide, including L-lactide, D-lactide, and polylactides based on both L-lactide and D-lactide. The ratio of L- and D-form content can vary. Typically, the PLA used in the method of the present invention contains both L- and D-forms of lactide. With respect to polylactide, the presence of D-form in the polymer usually affects the time it takes for the polymer to degrade. The content of D-form shortens the degradation time. In this way, PLA with a desired degradation time can be designed by varying the content of L- and D-form in the polymer. However, the content of D-form also affects the biomechanical properties of the resulting scaffold. Thus, the higher the content of D-form in the scaffold, the weaker the biomechanical properties. Therefore, it is important to select the appropriate balance between L- and D-form of polylactide for use in the method of the present invention. In general, the content of the D-form is typically about 1 to about 50% w / w, for example, about 2 to about 40% w / w, about 3 to about 35% w / w, or about 4 to about 30% w / w. In one embodiment of the present invention, the polylactide contains 96% of the L-form and 4% of the D-form; such a lactide is also called 96 / 4 poly(L / D) lactide.

[0036] In the context of the present invention, the term PLA includes polylactides or copolymers, such as those formed between lactide and glycolide, poly(lactide-co-glycolide) (PLGA), where the lactide and glycolide content may vary.

[0037] The raw PLA material used to obtain the PLA fiber mesh for use in the method of the present invention has an intrinsic viscosity - determined at room temperature - of about 1.5 dl / g to about 5 dl / g, such as about 1.5 dl / g to about 4 dl / g, about 1.7 dl / g to about 3 dl / g, about 1.7 dl / g to about 2.5 dl / g, or about 1.8 dl / g, about 1.9 dl / g, about 2.0 dl / g or about 2.1 dl / g, and a maximum monomer content of about 0% to about 1% by weight, such as about 0.1% to a maximum of 1% by weight.

[0038] Below, a more detailed description is provided regarding biodegradable PLA polymers suitable for use in the methods of the present invention.

[0039] Suitable polymers are biodegradable polymers. Suitable polymers are natural or synthetic polymers. In general, synthetic bioabsorbable polymers have been widely studied as tissue engineering scaffolds. The controllable chemistry and properties of synthetic bioabsorbable polymers, as well as the feature of being easily reproducible, go beyond their use as scaffold materials. Synthetic bioabsorbable polymers can be classified into various subgroups, such as esters, orthoesters, anhydrides, carbonates, and amides, depending on the susceptibility of the functional groups to hydrolysis. In particular, polyesters are used in several clinical applications due to the ease of degradation by hydrolysis of the ester bonds. Also, the degradation products of polyesters may be absorbed via metabolic pathways, and by adjusting their structure, they have the ability to change the rate of degradation.

[0040] Some of the most widely and earliest studied synthetic bioabsorbable polymers used in tissue engineering are polyesters. The uniqueness of poly(α-esters) lies in their vast diversity and synthetic versatility. Within the poly(α-ester) class, the poly(α-hydroxy acids), including the stereoisomeric forms of polyglycolide (PGA) and polylactide (PLA), are the most widely studied polymers.

[0041] In the method of the present invention, PLA is used. PLA is a thermoplastic biodegradable polymer produced by condensation polymerization from lactic acid derived from fermentation of sugars from carbohydrate sources such as corn, sugarcane and tapioca, or by ring-opening polymerization from lactide, a cyclic dimer of lactic acid. Due to its chiral carbon atom, lactic acid exists in two enantiomeric forms called L-lactic acid (S) and D-lactic acid (R), which occur in the metabolism of all animals and microorganisms. Condensation polymerization usually results in only low molecular weight PLA. High molecular weight PLA can be obtained by ring-opening polymerization, which involves polycondensation of lactic acid followed by depolymerization to the dehydrated cyclic dimer lactide. Optically active lactide can be found as either D-lactide, L-lactide or meso-lactide (D,L-lactide). In addition to the three diastereomeric structures, racemic lactide, a racemic mixture of D-lactide and L-lactide also exists. The structure and composition of the polymer chains, and especially the ratio of the L-isomer to the D-isomer of lactic acid, affect the processing, crystallization and degradation behavior of PLA. By using copolymerization of L-lactide with meso-L,D- or racemic lactide, high molecular weight amorphous or semi-crystalline polymers with a melting range of 130-185 °C can be obtained. Poly(L-lactide) (PLLA), a homopolymer containing only L-lactide, is semi-crystalline and has the highest melting point, whereas PLA copolymers with higher D-isomer content have lower melting points and dramatically lower crystallization behavior, becoming amorphous when the D content is higher than 12-15%.

[0042] Since PLA is an aliphatic polyester, it is susceptible to hydrolysis due to the presence of ester groups in its structure. The hydrolysis behavior, rate and mechanism can be controlled by changes in molecular structure and higher order structures, as well as by factors of the medium, such as temperature, pH and catalytic species (e.g., alkali and enzymes) of PLA. The in vivo hydrolysis rate is comparable to the in vitro degradation, so the in vitro degradation behavior and rate can predict the in vivo degradation to some extent. PLA does not require the presence of enzymes to catalyze hydrolysis. Lactic acid is produced in the metabolism of living organisms, and as a result, the degradation products of PLA are non-toxic. The hydrolysis of aliphatic polyesters begins with the uptake of water into the matrix, followed by hydrolytic cleavage of the ester bonds. The initial crystallinity affects the hydrolytic degradation rate, since the amorphous part has a higher rate of water uptake and the crystalline segments reduce the water permeability in the matrix. The autocatalytic effect of PLA specimens has also been reported. Autocatalysis is due to the increase in the number of compounds containing carboxy end groups in the center of the specimen when compounds with low molar mass cannot penetrate the outer shell where the degradation products dissolve in the surrounding solution.

[0043] PLA can be processed into various forms due to its thermoplastic nature. Melt processing is the most widely used method for PLA. In addition, injection molding and extrusion are widely used methods to produce PLA films and fibers for different nonwovens or fibers. Electrospinning of PLA is also used in medical applications to produce thin fibers that can be used as medical tissue scaffolds, wound dressings, drug carriers, protective fabrics and nanocomposite materials. The wide range of medical applications of PLA includes orthopedic screws, tissue engineering scaffolds, sutures, protein encapsulation and delivery, microspheres and drug delivery systems. PLLA is a slow degrading polymer (taking 2-5 years or more to be fully resorbed in vivo) with good tensile strength, low elongation and high modulus. Therefore, PLLA is considered ideal for load-bearing applications such as orthopedic fixation devices. PDLLA, on the other hand, degrades faster, losing its strength within 1-2 months and, when hydrolyzed, loses mass within 12-16 months. PDLLA also has a lower tensile strength compared to PLLA. Therefore, PDLLA is preferred as a drug delivery vehicle and as a low-strength scaffold material for tissue engineering. PLLA (semi-crystalline), PLDLA (amorphous), P(L / DL)LA70 / 30 (amorphous) and P(L / D)LA96 / 4 (semi-crystalline) are the most commonly used PLA polymers in the medical industry.

[0044] Polylactide copolymer PLGA is the most studied degradable polymer for biomedical applications. Its properties are significantly different from PLA and PGA, so different copolymer compositions can optimize PLGA for different applications. At 25-75% lactide composition, PLGA forms an amorphous polymer that is highly hydrolytically unstable compared to more stable homopolymers. Several different processing techniques have been used to fabricate PLGA scaffolds, such as gas foaming, microsphere sintering, porogen leaching, electrospinning and polymer printing. PLGA has been used specifically as sutures and drug delivery devices, as it degrades rapidly compared to other polyesters. PLGA has also been fabricated into tissue engineering scaffolds, as it exhibits excellent cell attachment and proliferation properties.

[0045] As mentioned above, in the context of the present invention, the term PLA includes all polylactide polymers (also called polylactic acid polymers) and copolymers with lactide and glycolide. In certain embodiments, the PLA is a polylactide polymer.

[0046] With regard to stability, PLA contains ester bonds that can degrade into lower molecular weight PLA, or monomers, dimers, etc. In the method of the invention, measures are taken to avoid extensive degradation. The rate of degradation is temperature and pH dependent. In the method of the invention, after melt spinning, at least some of the or all method steps are carried out at most room temperature, i.e. at most 25-30°C, for example at most 25°C, and a few of the method steps (e.g. some steps of drying and sterilization) are carried out at temperatures clearly below room temperature, for example below 0°C, for example below -10°C, -20°C or -25°C.

[0047] The stability of the PLA component of the PLA-collagen scaffold is ensured after processing by changes in the inherent viscosity or monomer content (see, e.g., Example 1 herein). The overall stability of the PLA-collagen scaffold is ensured after all processing steps by, e.g., biomechanical characterization, as shown in Examples 2 and 3 herein.

[0048] definition As used herein, "three-dimensional material" or "three-dimensional structure" refers to any material that has height, width and depth. One example of a three-dimensional structure is a scaffold. The three-dimensional material of the present invention is preferably implantable, biodegradable and biocompatible.

[0049] As defined herein, a "biodegradable material" is a material that, after introduction into the body, breaks down into soluble and non-toxic by-products, thereby eliminating the need for removal or further manipulation.

[0050] As defined herein, an "implantable material" is a material of any shape or size that is suitable for implantation into a subject.

[0051] As defined herein, a "biocompatible material" is a material that is not harmful or toxic to living tissue.

[0052] As defined herein, "loading", e.g., loading a fiber mesh with collagen, is intended to mean a process by which collagen is added to the fibers or contacted with a fiber mesh such that the collagen is found on the fiber mesh, incorporated into the fiber mesh, or both, or the fiber mesh is impregnated with collagen.

[0053] The term "biomechanical strength" as used herein is intended to refer to the ability of a scaffold to be applied to biological tissue without breaking into fragments while maintaining its functionality, and to withstand normal handling of the scaffold after manufacture, during storage, and during application.

[0054] The term "mechanical strength" as used herein is intended to refer to the ability of a scaffold to withstand normal handling of the scaffold after manufacture, during storage, and during application. In this context, "mechanical strength" is sometimes used synonymously with "biomechanical strength" and "biomechanical function."

[0055] In the context of the present invention, a desired improvement in biomechanical properties means i) an increase in one or more biomechanical parameters or biomechanical characteristics, ii) a decrease in one or more biomechanical parameters or biomechanical characteristics, or iii) no change in one or more biomechanical parameters or biomechanical characteristics. The improvement is based on the measurement of the biomechanical characteristics or biomechanical parameters of a scaffold prepared according to the present invention compared to a scaffold prepared by the same method, but omitting the final step of sterilization. A desired improvement is when the increase (or decrease) is 1% or more, 2% or more, 3% or more, 4% or more, 5% or more, 6% or more, 7% or more, 8% or more, 9% or more, or 10% or more; or when there is no change of less than 10%, less than 9%, less than 8%, less than 7%, less than 6%, less than 5%, less than 4%, less than 3%, less than 2%, or less than 1%. The data used are average values ​​based on two or more measurements.

[0056] In the context of the present invention, relevant biomechanical characteristics or parameters include the following: invariant modulus, invariant creep modulus (i.e., invariant modulus under creep conditions), dynamic invariant modulus, memory value, fluid mobility, apparent permeability, permeability in creep (i.e., apparent permeability under creep conditions), dynamic modulus (i.e., stress / strain ratio), and stiffness (stress / strain ratio of the material). The parameters are further defined in the experimental section herein.

[0057] In the context of the present invention, the term "creep test" is explained as follows: Creep tests are pseudo-static in nature (time change in strain under constantly applied stress) and are often used to evaluate the viscoelastic properties of materials and to approximate them in some models.

[0058] As used herein, the term "fibrous" refers to a material consisting of fibers. Fibers with diameters of only one size or different sizes can be used in the preparation of meshes for use in the method of the invention. These polymer fibers are selected from PLA fibers with diameters of 5-100 μm, for example 5-75 μm, 5-50 μm, 5-40 μm, 5-35 μm, 10-75 μm, 10-50 μm, 10-40 μm, 10-35 μm, 15-75 μm, 15-50 μm, 15-40 μm, or 15-35 μm. The diameter is the average diameter of the fibers in the structure. The cross section of the fibers may be round, but also any other shape, such as elliptical, star-shaped, rectangular, triangular, etc.

[0059] In the context of the present invention, "porosity" is calculated as follows: porosity = pore volume / specimen volume x 100%. Similarly, porosity is considered to be evaluated by micro-computed tomography (microCT), which can analyze the overall 3D structure of the structure.

[0060] In the context of the present invention, a "bioactive" is a substance, compound and / or biological material that has biological or pharmacological activity, i.e., the activity is an effect on an organism, tissue and / or cell, such as, but not limited to, the beneficial or adverse action of a drug on an organism. When drugs are complex chemical mixtures, this activity is exerted by the active ingredient or pharmacophore of the substance, but may be modified by other components. Typical examples are antibiotics, enzymes and vitamins, implants and cells.

[0061] In the context of the present invention, the term "about" is intended to indicate a range corresponding to a range from the stated value -10% to the stated value +10%.

[0062] Method for obtaining fibers and fiber meshes Fibers of PLA polymers can be obtained by various methods well known to those skilled in the art. Fibers of PLA polymers can be obtained by melt processing or electrospinning. As is evident from the examples herein, melt spinning has proven suitable for preparing PLA fibers for use in the method of the present invention.

[0063] The melt spinning process involves melting a polymer or heating the polymer into a soft form. Thus, depending on the selection of a particular polymer, a suitable temperature is selected for the spinning process. Typically, the temperature ranges from about 60 to about 300°C, such as from about 70 to about 250°C when polylactide is used.

[0064] Polylactides can be either amorphous or semi-crystalline, depending on the ratio of L-lactide and D-lactide monomers. Polyglycolide is semi-crystalline. PLA is a brittle polymer with a melting point range of about 170-180°C and a glass transition temperature of about 63°C. The glass transition temperature of P(L / D)LA is about 60°C, and semi-crystalline poly(L,D-lactide) has a melting range of about 135-170°C. Amorphous polymers have no melting point. Polyglycolide is produced by ring-opening polymerization of glycolide. Polyglycolide has a crystallinity of about 45-55%. Polyglycolide has a high melting point (about 225°C), and a glass transition temperature of about 35°C. Polyglycolide decomposes relatively quickly into acidic products.

[0065] Generally, the melt spinning process of PLA is carried out at temperatures ranging from about 60°C to 300°C.

[0066] The PLA raw material is dried before spinning, and protective gas is used to prevent decomposition during the spinning process.

[0067] The melt spinning process results in PLA fibers suitable for use in the PLA fiber mesh for the present invention. Although some changes may occur in the PLA material during the spinning process, these changes should not significantly affect the suitability of the resulting fiber mesh to be used to obtain a PLA-collagen scaffold. The resulting fibers are semi-crystalline. Generally, the resulting fibers have a monomer content close to the initial monomer amount of the PLA and / or have a decrease in intrinsic viscosity of less than 30%, 25% or 20% compared to the feedstock. This is especially true when the intrinsic viscosity of the PLA feedstock is up to 2.5 dl / g. However, when the PLA feedstock has an intrinsic viscosity of 2.5 to about 5 dl / g, the resulting fibers may have a decrease in intrinsic viscosity of about 70% or less, such as about 60% or less, about 50% or less, about 40% or less, about 30% or less, about 25% or less, or about 20% or less. Generally, suitable scaffolds are obtained when the fibers resulting from the spinning process have an intrinsic viscosity in the range of 1.5 to about 5 dl / g, such as 1.5 to 4 dl / g, or about 1.5 to about 3 dl / g, where the intrinsic viscosity is determined as described herein at room temperature.

[0068] Once the fibers are obtained, they are transformed into a mesh. As mentioned above, the mesh is carded to form a 3D network. Carding is a mechanical process that disentangles and mixes the fibers, producing a continuous, randomly oriented web, i.e. a carded mesh or fiber network. Carding breaks the staple locks and unorganized clumps of fibers, then aligns the individual fibers so that they are almost separate from each other. To obtain the desired porosity of the network thus obtained, the network is subjected to needle punching, a process that uses needles with notches along the needle shaft that grab the top layer of fibers as the needle enters the fiber mesh / mat and intertwines with the inner layer of fibers. Needle punching creates an intertwined and compressed 3D network from the card, improving the mechanical properties, but still leaving the structure very porous.

[0069] The fiber mesh, or fiber 3D network thus obtained, has a porosity of at least 85%. Exemplary porosities are 85%, 86%, 87%, 88%, 89%, 90%, 91%, 92%, 93%, 94%, 95%, 96%, 97%, 98% and 99%. Thus, the 3D network can have a porosity of 85-99%. As defined herein, porosity, i.e., porosity, is a measure of the interstitial (i.e., "empty") space in a material, and is the percentage of the volume of interstitial spaces over the total volume. The resulting mesh or 3D network has a porous structure, and in particular a pore network throughout the material.

[0070] The fiber mesh, or 3D network, can be cut to a desired shape and / or size.

[0071] The thickness of the resulting 3D network is typically between 0.1 and 50 mm.

[0072] Filling of fiber mesh with collagen Before loading the fiber mesh or 3D network with collagen, the fiber mesh or 3D network may be subjected to a washing procedure. The washing procedure is typically carried out in an aqueous medium, an alcoholic aqueous medium, or in an alcohol, such as ethanol. The washing procedure may be carried out several times, followed by drying of the mesh or 3D network. Drying can be carried out by freeze-drying or by drying at room temperature, for example in a laminar hood. The fiber mesh or 3D network can also be cut into pieces of a size suitable for the intended use.

[0073] The resulting fiber mesh or 3D network is filled with collagen. Filling may be performed mainly to ensure filling of the outer surface of the mesh or 3D network, or to load collagen into the gaps of the mesh or 3D network. Collagen is also adhered to the fibers of the mesh or 3D network. Collagen may be present on the surface, for example, by adhesion, and may be present in the mesh or 3D network by combination, for example, collagen may be present in the gaps by capillary forces and / or on the surface by adhesion. It is assumed that the mechanism by which collagen is included in the scaffold is not of primary interest, but in the context of the present invention, collagen is intended to be present in the gaps and on the surface of the mesh or 3D network.

[0074] Typically, loading with collagen involves dissolving or dispersing the collagen in a suitable medium. Generally, the medium is an aqueous medium. The pH and / or viscosity of the aqueous medium containing the collagen is adjusted.

[0075] The concentration of collagen in a suitable medium is about 0.1% to about 5% w / w, for example, about 0.1% to about 4% w / w, 0.1% to about 3% w / w, or about 0.1% to about 2% w / w.

[0076] The mesh or 3D network may then be immersed in a collagen-containing medium. Alternatively, the mesh or 3D network is immersed in a solvent (that does not dissolve the scaffold) and a solution / dispersion of collagen is added to the immersed mesh or 3D network. Collagen is also loaded into the mesh or 3D network by injecting or spraying the mesh or 3D network with a collagen-containing medium.

[0077] The pH of the collagen aqueous medium should be kept below 8 during processing. The process is preferably carried out at room temperature (RT), but may be carried out at elevated temperatures not exceeding 45°C. The weight ratio of PLA:collagen is about 95% to about 60% by weight of PLA and 40% to about 5% by weight of collagen, and preferably about 90% to 75% by weight of PLA and 25% to 10% by weight of collagen. Thus, the weight ratio between PLA and collagen (PLA / collagen) is 60 / 40 to 95 / 5, such as 70 / 30 to 90 / 10, 75 / 25 to 80 / 20, or 80 / 20 to 90 / 10. The porosity of the PLA-collagen scaffold, before possible loading with cells, drugs, etc., is about 70% to 99%, preferably about 80% to 95%.

[0078] As defined herein, recombinant human collagen refers to a human collagen polypeptide produced by using recombinant techniques, e.g., by using appropriate polynucleotides, expression vectors and host cells. Recombinant techniques are well known to those skilled in the art, e.g., several commercially available recombinant human collagens are available on the market.

[0079] The use of recombinant human collagen reduces the risk of transmitting known and unknown animal-derived pathogens and undesirable immunological responses. Moreover, unlike other naturally derived materials for cartilage regeneration, recombinant human collagen does not suffer from batch-to-batch variability. Therefore, recombinant human collagen can be produced in large quantities and with uniform quality at the grade required by Good Manufacturing Practice (GMP).

[0080] Collagen is the most abundant protein in the extracellular matrix (ECM) and is the main component of skin and musculoskeletal tissues. At least 28 different types of collagen have been identified to date. All collagens have a triple helical structure in which three individual chains, each in a left-handed polyproline II helix, are coiled together to form a right-handed supercoiled triple helix with a rope-like structure. In this structure, collagens exhibit a characteristic repeating sequence, glycine-XY, where glycine is an amino acid small enough to pack into the center of the triple helix structure. The positions of X and Y can be any amino acid, but X is often proline and Y is often hydroxyproline. Collagens can be divided into different groups based on their structure and supramolecular organization. These groups include fibril-forming collagens, fibril-associated collagens (FACIT), network-forming collagens, anchoring fibrils, transmembrane collagens, and basement membrane collagens, each of which has its own unique functions. The most abundant group of collagens are the fibril-forming collagens, which account for about 90% of all collagen. Type I collagen is the most abundant and best studied collagen, forming over 90% of the organic mass of bone and is the major collagen of tendons, skin, ligaments, cornea and many interstitial connective tissues, except for a few tissues such as hyaline cartilage, brain and vitreous body. Type II collagen, on the other hand, is the characteristic and major component of hyaline cartilage, but is also found in the vitreous body, corneal epithelium, notochord, nucleus pulposus of the intervertebral disc, and at the embryonic epithelial-mesenchymal transition.

[0081] Collagen has high mechanical strength, good biocompatibility, low antigenicity, and crosslinking ability, allowing for tuning of collagen's mechanical, degradation, and water uptake properties. When fabricated into highly porous scaffolds, crosslinking of collagen is necessary to produce scaffolds with sufficient mechanical properties and degradation rates. Collagen has been widely studied for use in various medical applications, as well as a wide range of tissue engineering applications, such as bone, cartilage, and skin tissue engineering. Modifying or combining collagen with other degradable polymers improves collagen's potential as a biomaterial. The major collagen that can be easily prepared in pure form and in commercial quantities is type I collagen. Type I collagen is the most widely used collagen for tissue engineering applications. Collagen scaffolds offer an alternative means to provide biological information to growing constructs, unlike biodegradable synthetic polymers. A wide range of cell adhesion and other signals are achieved while using collagen scaffolds, improving the quality of tissue engineering products. Generally, freeze-drying or stereolithography is used to fabricate porous collagen-based scaffolds, and carbodiimides are often applied for crosslinking. Porous collagen scaffolds are often combined with other components such as bioceramics or synthetic biodegradable polymers. Collagen is already available in various commercial medical products, such as bioprosthetic heart valves or wound dressings. The use of collagen can be divided into two different categories: tissue-based devices (bioplastic heart valves) that use natural stabilizing tissue as the device, and purified collagen, where the collagen is solubilized by an enzymatic digestion step and reconstituted into various products (wound dressings). Recombinant collagens have emerged because they offer a means to produce high purity, disease-free collagens and have the potential to produce all types of collagen, even those with very low abundance in natural tissues.

[0082] Suitable collagen for use in the method of the present invention is collagen from any source, including animal-derived, human, recombinant or synthetic collagen. Collagen may be collagen type I, II, III, IV, V, VI, IX or XI. Any combination of these collagen types may also be utilized. Collagen is more particularly collagen type I, II or III. Collagen may also be a combination of at least type I, type II and type III collagen, at least type I and type III collagen, at least type I and type II collagen, or at least type II and type III collagen. Any collagen may be used in combination with animal-derived, recombinant or synthetic collagen, or collagen may be used alone or in combination with other collagens.

[0083] As used herein, "recombinant human collagen material" refers to any material (eg, any solution or gel) that contains recombinant human collagen.

[0084] Collagen for use in the method of the invention may be porous. For example, freeze-drying makes collagen porous and elastic, and therefore well suited for purposes such as supporting chondrocyte proliferation and cartilage matrix production. Collagen, such as freeze-dried collagen networks, is an excellent microenvironment for cell attachment. In an embodiment of the invention, collagen is freeze-dried. Collagen may be freeze-dried as is (e.g., in the form of a collagen solution). The pore size of the collagen structure varies between 20 and 250 μm and can be selected from 20 to 250 μm, 50 to 250 μm, 30 to 200 μm, 40 to 200 μm, 50 to 200 μm, or 60 to 200 μm. It is also possible to convert collagen into a gel before freeze-drying, i.e., collagen may be in the form of a freeze-dried gel.

[0085] Collagen is used in scaffolds, whether as sponges, sheets or gels, that can be used for tissue repair and regeneration. Collagen scaffolds are believed to have the right properties to allow tissue regeneration, such as pore structure, permeability, and hydrophilicity. Collagen scaffolds support cell attachment, robust cell spreading, proliferation, and 3D differentiation.

[0086] The PLA-collagen scaffold thus obtained is then dried. Drying may involve freeze-drying, possibly with a preceding step of freezing the wet scaffold at -20 to -50°C for more than 10 hours. Freezing depends on the material used and the size of the specimen, and usually takes more than 10 hours. The freezing temperature affects the size of the formed pores in the structure and also depends on the material used. The freezing temperature generally varies between -10°C and -80°C, with -20°C to -50°C being preferred. The freeze-drying process is usually carried out in the same temperature range as the freezing process or at a lower temperature. The freeze-drying time depends on the material used and the size of the specimen, and usually varies between 24 and 48 hours.

[0087] After filling the mesh or 3D network with one or more collagens, a PLA-collagen scaffold, optionally dried, is obtained and preferably subjected to cross-linking of the collagens of the PLA-collagen scaffold, which is performed to increase the biomechanical strength of the final scaffold and make the collagen component more stable in vivo.

[0088] Suitable cross-linking methods are well known to those of skill in the art and include, but are not limited to, the use of chemical cross-linking agents such as 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC), glutaraldehyde, genipin and UV light, or combinations thereof.

[0089] A suitable combination of crosslinkers is EDC, preferably EDC is used together with N-hydroxysuccinimide (NHS). Crosslinking is typically carried out in an alcohol medium, such as in an ethanol-water solvent. The ethanol content is usually 50% to 99% by weight, for example 60% to 99% by weight, 70% to 99% by weight, 80% to 99% by weight, or 90% to 99% by weight.

[0090] Crosslinking is typically performed by using EDC and NHS. The concentration of EDC may range from 5 to 30 mM, such as 5 to 25 mM, 10 to 20 mM, such as 14 mM EDC. The concentration of NHS may range from 1 to 10 mM, 5 to 10 mM, or about 6 mM NHS.

[0091] The crosslinked PLA-collagen scaffold is typically washed extensively with water, an alcoholic aqueous medium, or an aqueous medium containing alcohol to avoid crosslinking residues.

[0092] The crosslinked PLA-collagen scaffold is then dried. Drying may include lyophilization, optionally with a prior step of freezing the wet scaffold at -20 to -50°C for more than 10 hours. Details regarding lyophilization and freezing are similar to those described above. Temperature can affect the size of the pores in the final product; thus, the lower the temperature, the smaller the pore size.

[0093] Sterilization of dry PLA collagen scaffolds In addition to being safe to use and functioning properly in, for example, an AC environment, the scaffold should be sterile prior to implantation. One of the most reliable methods of sterilization of implantable medical devices is gamma irradiation. Gamma irradiation is highly effective and free of residual chemicals that can cause cytotoxicity. However, gamma irradiation is known to affect the properties of biodegradable polymers such as PLA and collagen. Collagen in particular may suffer loss of mechanical integrity during too high levels of irradiation. Collagen is temperature sensitive and elevated temperatures during processing and sterilization should be avoided.

[0094] After drying of the PLA-collagen scaffold, it is sterilized, resulting in a sterilized PLA-collagen scaffold. Any suitable method can be used that does not result in undesired degradation of the PLA, collagen or PLA-collagen scaffold. A suitable method is to subject the PLA-collagen scaffold to gamma irradiation, but as mentioned herein, any temperature increase should be avoided to avoid any degradation of the collagen component. Since irradiation is a process involving the transfer of energy, it should be carried out under temperature-controlled conditions for sensitive materials such as collagen. A method to keep the temperature low is to cool the PLA-collagen scaffold during the irradiation process. The scaffold is typically cooled to a temperature of about -200°C to about 25°C before sterilization is performed. This method keeps the temperature below 40°C during sterilization.

[0095] The temperature of the PLA-collagen scaffold before sterilization is essentially the same as or higher than the temperature of the PLA-collagen scaffold during sterilization.

[0096] Sterilization is typically carried out at a temperature in the range of -200°C to 40°C, such as up to 30°C or up to 25°C, or at a temperature in the range of -100°C to 25°C, such as -70°C, -40°C, 0°C, 10°C, 20°C or 25°C.

[0097] As will be apparent from the examples herein, the doses used, when sterilization is by gamma irradiation, are typically in the range of from 10 kGy to about 27 kGy, for example 15-26 kGy, 16-25 kGy, such as 18 kGy, 19 kGy, 20 kGy, 21 kGy, 22 kGy, 23 kGy, 24 kGy, 25 kGy, 26 kGy, or 27 kGy.

[0098] Typically, the PLA-collagen scaffold is packaged in a suitable package prior to sterilization to ensure that the scaffold maintains its sterility after the sterilization process.

[0099] It is noted that sterilization can be carried out at any step of the method of the invention: thus, the fiber mesh (PLA mesh) can be subjected to sterilization before being filled with collagen, and the collagen itself can be sterilized before being filled into the PLA fiber mesh or the 3D network.

[0100] The main advantage of employing a sterilization process is that the resulting sterile scaffolds have fairly uniform biomechanical properties. As seen in Example 3, non-sterile but aseptically manufactured scaffolds have much more diverse biomechanical properties. From a regulatory point of view, it is a great advantage to be able to prepare scaffolds without significant batch-to-batch variation in properties. The reason is that the biomechanical properties give an indication of how the scaffold will behave in vivo and can predict the in vivo behavior, i.e., no or only slight variations should only slightly affect the in vivo properties, whereas larger variations may affect the in vivo properties and should be investigated whether such effects negatively affect the efficacy of the scaffold (e.g., scaffold stability, ability to remain at the site of administration, negative effects on the release of cells or bioactives, undesirable side effects, etc.). Therefore, strict control of obtaining identical or nearly identical biomechanical properties is also desired in relation to regulatory approval. In general, a variation in properties of ±10% can usually be tolerated.

[0101] As can be seen from the examples herein, scaffolds produced according to the present invention have improved biomechanical properties compared to non-sterile scaffolds.

[0102] The improved biomechanical properties may be expressed as a change in one or more biomechanical parameters or biomechanical properties, such as the constant modulus, constant creep modulus (i.e., constant modulus under creep conditions), dynamic constant modulus, memory value, fluid mobility, apparent permeability, permeability in creep (apparent permeability under creep conditions), dynamic modulus (stress / strain ratio), stiffness (i.e., stress / strain ratio of the material), etc. The desired change may be i) an increase, ii) a decrease, or iii) no change.

[0103] In one embodiment, the one or more biomechanical parameters or characteristics are selected from an unchanged creep modulus when tested under wet conditions and a dynamic modulus when tested under dry conditions. The desired change compared to the non-sterile scaffold is an increase. Typically, the increase is 1% or more, 2% or more, 3% or more, 4% or more, 5% or more, 6% or more, 7% or more, 8% or more, 9% or more, or 10% or more.

[0104] Some biomechanical parameters may decrease or remain unchanged. Such parameters include biomechanical parameters or biomechanical characteristics selected from creep modulus and dynamic modulus permeability, both determined under wet conditions. Desired decreases are 1% or more, 2% or more, 3% or more, 4% or more, 5% or more, 6% or more, 7% or more, 8% or more, 9% or more, or 10% or more, while no change is less than 10%, less than 9%, less than 8%, less than 7%, less than 6%, less than 5%, less than 4%, less than 3%, less than 2%, or less than 1%.

[0105] In particular, the improved biomechanical property is improved stiffness.

[0106] All details concerning the individual steps mentioned in the main aspect of the invention apply mutatis mutandis to all other aspects of the invention and are therefore not included in the following paragraphs.

[0107] Other aspects of the invention The present invention provides a method for preparing a sterile PLA-collagen scaffold, the method comprising: a) providing PLA in solid form; b) subjecting the PLA to a process, thereby obtaining fibers of PLA; and c) subjecting the obtained fibers to a process, thereby obtaining a mesh of fibers; d) optionally subjecting the mesh of fibers to carding and / or needle punching to obtain a 3D network of PLA fibers; e) optionally subjecting the mesh of fibers or the 3D network of fibers to one or more washing procedures; f) providing collagen in the form of a solution or gel; g) immersing the mesh of fibers or the 3D network of fibers obtained after step c) or optionally after steps d) or e) in a collagen solution or collagen gel to obtain a PLA-collagen scaffold; h) optionally drying the PLA collagen scaffold; and i) sterilizing the PLA-collagen scaffold obtained in step g) or optionally in step h).

[0108] The present invention provides a method for preparing a sterile PLA-collagen scaffold, comprising: a) providing PLA in solid form; b) subjecting the PLA to a process, thereby obtaining fibers of PLA; c) subjecting the obtained fibers to a process, thereby obtaining a mesh of fibers; d) subjecting the mesh of fibers to carding and / or needle punching to obtain a 3D network of PLA fibers; e) subjecting the mesh of fibers or the 3D network of fibers to one or more washing procedures; f) providing collagen in the form of a solution or gel; g) immersing the mesh of fibers or the 3D network of fibers obtained after step d) or optionally after step e) in a collagen solution or collagen gel to obtain a PLA-collagen scaffold; h) optionally drying the PLA collagen scaffold; and i) sterilizing the PLA-collagen scaffold obtained in step g) or, optionally, in step h).

[0109] The present invention provides a method for preparing a sterile PLA-collagen scaffold, comprising: a) providing PLA in solid form; b) subjecting the PLA to a process, thereby obtaining fibers of PLA; c) subjecting the obtained fibers to a process, thereby obtaining a mesh of fibers; d) subjecting the mesh of fibers to carding and / or needle punching to obtain a 3D network of PLA fibers; e) subjecting the 3D network of fibers to one or more washing procedures; f) providing collagen in the form of a solution or gel; g) immersing the 3D network of fibers obtained after step e) in a collagen solution or collagen gel to obtain a PLA-collagen scaffold; h) optionally drying the PLA collagen scaffold; and i) sterilizing the PLA-collagen scaffold obtained in step g) or optionally in step h).

[0110] The present invention provides a method for preparing a sterile PLA-collagen scaffold, comprising: a) providing PLA in solid form; b) subjecting the PLA to a process, thereby obtaining fibers of PLA; c) subjecting the obtained fibers to a process, thereby obtaining a mesh of fibers; d) subjecting the mesh of fibers to carding and / or needle punching to obtain a 3D network of PLA fibers; e) subjecting the 3D network of fibers to one or more washing procedures; f) providing collagen in the form of a solution or gel; g) immersing the 3D network of fibers obtained after step e) in a collagen solution or collagen gel to obtain a PLA-collagen scaffold; h) drying the PLA collagen scaffold; and i) sterilizing the PLA-collagen scaffold obtained in step h).

[0111] Uses of the scaffolds prepared by the method of the present invention The scaffolds obtained according to the present invention are typically used in human and veterinary surgical care. They can also be used in other health and medical care of humans and animals, i.e. companion animals, as well as in cosmetics.

[0112] The scaffold obtained according to the present invention can be used to treat articular surface pathologies, especially weight-bearing joints such as the knee. These pathologies present with symptoms such as pain and locking of the affected joint, requiring surgical intervention. Although these pathologies have different etiologies, cartilage damage can be the result of trauma, degenerative joint diseases such as osteoarthritis (OA), and developmental disorders such as osteochondritis dissecans.

[0113] The scaffold may be used as is, or may be loaded or combined with one or more of the following: tissue specific cells such as chondrocytes, somatic stem cells such as mesenchymal stromal cells from bone marrow or embryonic stem cells, cellular components such as growth factors or cytokines, blood components and fractions such as platelets or platelet rich plasma, or as drug substances, anti-inflammatory drugs such as ibuprofen. Of primary importance for the above conditions is the use of the scaffold of the present invention as a lesion filler, and of secondary importance the incorporation of potential additives.

[0114] The scaffolds obtained according to the present invention can also be used as dermal fillers in cosmetic surgery, e.g., facial enhancement.

[0115] Scaffolds are usually delivered surgically, i.e., implanted into the body. Scaffolds can be used for cartilage or osteochondral lesions. In the context of cartilage-related conditions, the scaffold is implanted into the debrided lesion site. The scaffold is secured to the lesion bed by absorbable sutures or fibrin glue. [Brief description of the drawings]

[0116] [Figure 1a] Graph of creep test, permanent modulus (a), memory value (b) and permeability (c) obtained under wet conditions (Example 3). [Figure 1b] Graph of creep test, permanent modulus (a), memory value (b) and permeability (c) obtained under wet conditions (Example 3). [Figure 1c] Graph of creep test, permanent modulus (a), memory value (b) and permeability (c) obtained under wet conditions (Example 3). [Diagram 2] 1 is a graph of dynamic (a) and static (b) gradient modulus of elasticity from a dynamic strain sweep test under wet conditions (Example 3). [Diagram 3] 1 is a graph of dynamic modulus vs. test temperature under dry conditions. As can be seen, there is no effect of temperature on modulus values, but there is a statistically significant difference (p=0.00) between non-sterile and sterilized samples (Example 3).

[0117] The invention is further illustrated in the following non-limiting examples. Working Example Example 1

[0118] A method for preparing a medical sterile scaffold with polylactide fibers and collagen, drying the structure by freeze-drying to realize the scaffold, and sterilizing the scaffold to obtain a sterile and stabilized PLA-collagen scaffold, rhCo-PLA. This example demonstrates that the PLA component is not negatively affected by the processing method.

[0119] The scaffolds were fabricated as follows: The PLA component, medical grade poly(L / D) lactide 96 / 4 (Corbion, Purac Biochem bv, Gorinchem, The Netherlands), with an intrinsic viscosity ranging from 1.8 dl / g to 2.2 dl / g and residual monomer content less than 0.1% was used to produce thin fibers by melt spinning. Prior to melt spinning, the PLA raw material was dried in a vacuum oven. Melt processing of the fibers was carried out by melt spinning under a protective atmosphere within the temperature range of 70–240 °C. The spinning apparatus consisted of a micro-extruder and a high-speed spinner. The fibers were cut into staple fibers and carded into mesh. PLA felt was produced by needle punching the carded PLA mesh. The PLA felt was washed and dried in a laminar hood and then packaged before being placed in a cleanroom environment. Fiber formation of recombinant human collagen type III (FibroGen Ltd., CA, USA) was performed by increasing the pH of the collagen solution to 7 using a basic buffer. After the formation of collagen gel, PLA felt was completely immersed in the collagen gel and placed into a sample mold. These structures were then freeze-dried to achieve completely dry structures, i.e., rhCo-PLA scaffolds. The fabricated scaffolds were cross-linked with a 95% ethanol solution containing 14 mM EDC (N-[3-dimethylaminopropyl]-N'-ethylcarbodiimide hydrochloride (Sigma-Aldrich, Helsinki, Finland) and 6 mM NHS (N-hydroxysuccinimide (Sigma-Aldrich, Helsinki, Finland) at room temperature (RT). The fabricated rhCo-PLA scaffolds were then washed, freeze-dried, packed, and then sent for sterilization by gamma irradiation ≤25 kGy under dry ice to avoid temperature rise during sterilization.

[0120] The following tests were performed on the prepared rhCo-PLA. The monomer content of L-lactide was measured by GC-MS method (Rambol Analytics, Lahti, Finland) with a lower limit of 0.01 wt%, and the intrinsic viscosity (iv) was measured with a Lauda PVS viscometer (Lauda DR.R.Wobster GmbH, KG, Königshofen, Germany). Samples were prepared by dissolving the polymer in chloroform at 1 mg / ml. Viscosity was determined using an Ubbelohde capillary viscometer type 0c (Schott-Gerate, Mainz, Germany).

[0121] By this manufacturing method, the sterilized medical scaffold, rhCo-PLA, made of polylactide fiber and collagen, achieved the following characteristics as shown in Table 1:

[0122] The residual monomer content of the PLA component did not change significantly, as the monomer content remained below 0.1 wt.%, the same as that of the raw material. The PLA processing temperature and sterilization used did not significantly change the intrinsic viscosity of the extruded PLA: an acceptable reduction in this experiment was about 50%.

[0123] [Table 1] Example 2

[0124] A medical sterile scaffold with polylactide fibers and collagen was produced by drying the construct by freeze-drying to realize the scaffold and sterilizing the scaffold to obtain a sterile PLA-collagen scaffold. The rhCo-PLA scaffold described in Example 1 was sterilized by gamma irradiation to evaluate the effect of sterilization, especially on the collagen component. Similarly, different gamma radiation doses (under dry ice) were used on the rhCo-PLA scaffold to show the effect of sterilization dose.

[0125] Data analysis was performed according to the procedures detailed in US Patent No. 10379106(B2). The analysis involved the extraction of invariant data such as viscous stiffness and memory values ​​under static and dynamic conditions, respectively, from the applied load (stress) and deformation (strain) data. In this specification, the following definitions were used:

[0126] An "invariant modulus" is a unique modulus value (ie, a true value) that is independent of time and frequency and can be used to predict the behavior of a material.

[0127] The "dynamic invariant modulus" is the ratio of the dynamic stress amplitude to the dynamic true (logarithmic) strain amplitude, expressed in real (not complex) algebraic terms (different from the commonly used definitions of real (storage) and imaginary (loss) moduli).

[0128] The "memory value" is a time-invariant property of a test specimen, with values ​​ranging between 0 and 1, which represents the viscous tendency of a material, even if the material itself is not fluid. The memory value has no theoretical prediction and must always be determined from experiments. In the present invention, it was experimentally measured separately for static (creep) and dynamic conditions, and it was found that the memory values ​​under these conditions are different.

[0129] "Fluid mobility" is a measure (coefficient) of the rate of fluid movement through a porous medium and is similar to the diffusion coefficient (mm 2 The fluid has the same units of velocity, per second. However, its nature differs from the latter because the fluid motion is due not only to diffusion but also to convection and momentum transfer. Fluid mobility describes how well a fluid can flow through itself under certain conditions.

[0130] "Apparent permeability" refers to the capacity of a porous medium to allow fluid to pass through its porous network and is expressed in squared distance units (m 2), which is the quantified topological ability of a material to transport a fluid through its porous structure and depends only on the material structure and not on the fluid properties. Here it differs from the permeability commonly defined by Darcy's law, the latter requiring an increase in a fluid pressure gradient across a specimen of material. However, as shown in US Pat. No. 10,379,106(B2), the method applied here also allows the measurement of permeability without knowledge of the fluid pressure gradient, and to distinguish between these methods, the term "apparent permeability" is used (expressed in millidarcy; 1 mDarcy = 10 -15 m 2 ).

[0131] Initial testing was performed on aseptically produced rhCo-PLA (rhCo-PLA-A) and compared it to rhCo-PLA sterilized at a standard irradiation dose of ≥25 kGy. In the process, the actual irradiation dose was measured to be 29 kGy (rhCo-PLA-S). Aseptically produced rhCo-PLA (rhCo-PLA-A) had a gamma-irradiated PLA component (sterilized at a standard irradiation dose of ≥25 kGy), but in aseptic conditions, the addition of the collagen component was done after sterilization, i.e., the collagen component itself was not irradiated. Therefore, the PLA fiber component should have identical properties and contribute to the overall biomechanical performance, and the differences were mainly due to the effect of collagen and its processing.

[0132] The fabricated scaffolds were then tested with a biomechanical testing procedure using dynamic mechanical analysis (DMA) in a standard compression sample holder (15 mm diameter) on a dynamic mechanical analyzer DMA242E (Netzsch Geratebau GmbH, Selb, Germany).

[0133] A portion of the scaffolds was subjected to creep testing under a constant stress of 0.2 N, while another portion was subjected to oscillatory force (strain-sweep method) at a frequency of 1 Hz resulting in strains ranging from 5 to 50 μm. Briefly, after establishing contact with the specimen on the probe and taring the offset, the starting height of the specimen immersed in the medium was measured again and further used as the starting height for the true strain calculations.

[0134] In all cases, the scaffolds were tested by fully immersing them in water at room temperature and allowing them to equilibrate for 15 min before measurements. Therefore, all specimens were fully impregnated, with no air bubbles or dry areas observed. The cross-sectional area of ​​the scaffolds tested ranged from 20 to 26 mm. 2 All tests were performed for up to 300 min (until the dimensional changes approached a constant value; displacement resolution ±0.0005 μm). These data were saved and exported as ASCII text files to a data processing software (Microsoft Excel supplemented with customized code). Primary data were converted to stress and true strain, and to the ratio of strain to stress versus experimental time. A numerical algorithm of time convolution was then applied, and the processed data were nonlocally integrated pairwise, row by row, with a mathematical method described in detail in US Pat. No. 10,379,106 (B2).

[0135] The experimental data are shown in Table 2.

[0136] [Table 2]

[0137] Table 2 shows that sterilization by standard procedure with gamma irradiation at a dose of ≥ 25 kGy affected the biomechanical properties of rhCo-PLA scaffolds as follows:

[0138] In the creep test, with a sterilization dose of 29 kGy, the rhCo-PLA-S scaffolds suffered a significant decrease in the immutable modulus (-22%) after irradiation. The memory value was also decreased (-43%), indicating that the rhCo-PLA scaffolds became more elastic after sterilization. The decrease in effective fluid mobility and apparent permeability of the sterilized rhCo-PLA scaffolds (almost doubled) indicates a state of reduced fluid mobility inside the scaffolds, and thus, that the irradiated scaffold structure became less permeable to fluids. These changes are undesirable, and such poor biomechanical properties of rhCo-PLA-S cannot be tolerated. Thus, the creep test showed that this sterilization method significantly affected the biomechanical properties of the rhCo-PLA scaffolds, making their use in the intended applications unwarrantable.

[0139] Strain sweep testing showed that the dynamic constant modulus increased (+33%), indicating that after sterilization, the rhCo-PLA-S scaffold became stiffer and less able to adapt to the dynamic strains of the surrounding tissue.

[0140] The second step was to evaluate the more precise gamma irradiation dose effect on rhCo-PLA scaffolds. Under the standard dose of gamma irradiation ≥ 25 kGy, a study was conducted to compare different doses for rhCo-PLA scaffolds. Thus, sterilization by low dose gamma irradiation for rhCo-PLA scaffolds was performed at the following doses: 18 kGy (G18), 20 kGy (G20), 22 kGy (G22), and 25 kGy (G25). Non-sterile rhCo-PLA scaffolds (G0) were used as a reference.

[0141] The fabricated scaffolds were tested with the biomechanical testing procedure using dynamic mechanical analysis, as previously described, using the same dynamic mechanical analyzer, DMA242E (Netzsch Geratebau GmbH, Selb, Germany). Scaffolds were subjected to creep testing under a constant force of 0.2 N as described above. In all cases, scaffolds were tested completely immersed in water at room temperature and allowed to equilibrate for 15 min before measurements. The area of ​​the tested scaffolds ranged from 37 to 46 mm. 2 It was.

[0142] The biomechanical testing results of the rhCo-PLA scaffolds with these different low doses of gamma irradiation are shown in Table 3. The results show that these different doses of gamma irradiation, ≦25 kGy, did not significantly change the biomechanics of the rhCo-PLA scaffolds (e.g., since the values ​​of the constant modulus and memory values ​​of the G0 sample are within the limits measured for the sterilized samples G18-G25).

[0143] [Table 3] Example 3

[0144] A medical sterile scaffold with polylactide fibers and collagen was fabricated by drying the construct by freeze-drying to realize the scaffold and sterilizing the scaffold to obtain a sterile PLA-collagen scaffold.

[0145] The rhCo-PLA scaffolds described in Example 1 were sterilized by gamma irradiation at RT (S-RT) or at low temperature (-70°C) (S-LT) to evaluate the effect of sterilization and compare it to a non-sterile scaffold (NS). This study demonstrated possible changes in the biomechanical properties of the scaffolds with different sterilization methods at a generally similar expected dose (25 kGy) and investigated whether lower temperatures during sterilization would affect the biomechanical properties. Biomechanical testing was performed on dry and wet samples as described below.

[0146] Biomechanical analysis was performed in dry conditions up to 60 °C and in wet soaked conditions at 25 °C in compression mode. The fabricated scaffolds were tested with a biomechanical testing procedure using dynamic mechanical analysis in a standard compression sample holder (15 mm diameter) on a dynamic mechanical analyzer DMA242E (Netzsch Geratebau GmbH, Selb, Germany). One part of the scaffolds was subjected to creep testing at a constant force of 0.2 N, while another part was subjected to strain sweeps in the amplitude range of 5 to 25 μm at a frequency of 1 Hz. For the latter, the loading cycle was repeated 10 times as in Example 2.

[0147] The scaffolds were tested completely immersed in water at 25°C and allowed to equilibrate for 15 min before measurements. Thus, all specimens were fully impregnated and no air bubbles or dry areas were observed. With the aim of evaluating the thermal stability of the mechanical properties, additional tests were performed on a set of dry specimens under dry conditions (air) at 1 Hz and less than 25 μm amplitude, but with heating to 60°C at a rate of 2 K / min.

[0148] The cross-sectional area of ​​the scaffolds tested was 30–40 mm 2The data analysis was performed according to the procedure described in US Pat. No. 10,379,106 (B2) and aimed to extract invariant data such as viscous stiffness and memory values ​​under static and dynamic conditions, respectively, using the data shown in Figures 1-3. It is worth mentioning that under repeated dynamic loading, all specimens contracted gradually in all loading sequence cycles. Therefore, the integral (slope) values ​​of the dynamic stress / strain ratio ("standard dynamic modulus") at 1 Hz and the associated static stress / strain ratio ("standard static modulus") were extracted to represent values ​​covering all loading cycles.

[0149] As can be seen from Figure 1, in creep tests, there are significant differences in some property values ​​between some specimen types, but not others (error bars = standard deviation at wet conditions). For example, S-RT specimens have lower creep coefficients and higher permeability in creep vs. S-LT specimens. Therefore, there is an effect of temperature control (reducing temperature) during the sterilization procedure.

[0150] [Table 4]

[0151] Under dynamic wetting conditions at 25 °C and 1 Hz (Fig. 2), there was no significant difference in the dynamic stress / strain ratio (slope modulus) and static stress / strain ratio (slope modulus) among the specimens (NS, S-RT, and S-LT).

[0152] Surprisingly, for the 1 Hz dynamic drying condition with temperature ramp up to 60 °C, there was a significant difference between the dynamic and static stress / strain ratio samples with p=0.00 (with a confidence interval of 95%, values ​​of p<0.05 are statistically significant for a two-tailed permutation t-test), indicating that the S-RT and S-LT scaffolds became stiffer after the sterilization process (Figure 3). These differences were between the non-sterile samples (NS) and the sterilized samples (S-RT and S-LT), but not between those sterilized at RT (S-RT) or at low temperature (S-LT). Temperature changes during dynamic testing did not affect the properties up to 60 °C. Also, since there appears to be no effect of the testing temperature on the stiffness of the samples (i.e., the stress / strain ratio of the material) up to 60 °C, it can be concluded that the observed differences are mainly due to the sterilization process and that the difference in temperature; room temperature (S-RT) or low temperature (S-LT) does not affect the properties of these samples.

[0153] These results indicate that sterilization (RT or low temperature (LT)) generally stabilizes the scaffold properties, but this is only evident in these tests in the dry state.

[0154] As shown in Figure 3 and Table 5, there is no effect of temperature on the modulus values, but there is a statistically significant difference (p=0.00) between non-sterile and sterilized scaffolds.

[0155] [Table 5]

[0156] Generally, it has been assumed that sterilization has a recessive effect on the material properties of scaffolds, but in this case sterilization leads to an unexpected result: gamma irradiation seems to have a positive effect on stabilizing the biomechanical properties of the scaffold.

[0157] In summary, this Example 3 demonstrates the following effects of the sterilization procedure on materials versus non-sterile materials (the changes are therefore classified as poor - change in property undesirable, moderate - neutral effect, no statistically significant difference, or good - change in property desirable):

[0158] [Table 6]

[0159] As stated above, a desired improvement means i) an increase in one or more biomechanical parameters or characteristics, ii) a decrease in one or more biomechanical parameters or characteristics, or iii) no change in one or more biomechanical parameters or characteristics.

[0160] As can be seen from the above table, an increase is desired for the following biomechanical parameters: unchanged creep modulus (wet conditions), dynamic modulus (dry conditions), and dynamic modulus versus temperature (dry conditions), while a decrease or no change is desired for the following biomechanical parameters: permeability at creep (wet conditions) and dynamic modulus (wet conditions).

[0161] As previously discussed herein, a desirable improvement is an increase (or decrease) of 1% or more, 2% or more, 3% or more, 4% or more, 5% or more, 6% or more, 7% or more, 8% or more, 9% or more or 10% or more; or no change of less than 10%, less than 9%, less than 8%, less than 7%, less than 6%, less than 5%, less than 4%, less than 3%, less than 2% or less than 1%.

[0162] Therefore, it can be assumed that the sterilization procedure of gamma irradiation at RT or low temperature (LT) is a positive step for these types of scaffolds to achieve a more stable structure in both dry and wet conditions in terms of kinetics. Similarly, sterilization at low temperatures gives favorable results compared to sterilization at RT, especially shown in the wet creep conditions.

[0163] It was also determined that the rhCo-PLA scaffold described in Example 1 had an unexpected responsiveness to gamma sterilization conditions (dose and temperature control), such as improved biomechanical stability, allowing for more precise control of mechanical properties and necessitating optimization of materials for clinical requirements. This effect was unexpected, since it is commonly known that irradiation weakens or even destroys many organic materials and polymers.

Claims

1. A method for preparing a sterile scaffold for medical use, comprising the steps of: i) filling a fiber mesh comprising fibers of polylactide polymer or copolymer (PLA) with collagen to obtain a PLA-collagen scaffold; ii) drying the PLA-collagen scaffold obtained from step i); iii) sterilizing the PLA-collagen scaffold obtained from the drying step ii) to obtain a sterile scaffold, wherein the sterilization is performed by gamma irradiation with a radiation dose of up to 25 kGy.

2. The resulting sterile scaffold has improved biomechanical properties compared to a non-sterile scaffold, the improved biomechanical properties being characterized by: a) Permanent Creep Modulus (Wet Conditions) b) Dynamic Modulus (dry conditions), and c) Dynamic Modulus vs. Temperature (Dry Conditions) The method of claim 1, wherein the increase is expressed as an increase of 1% or more in one or more biomechanical parameters or biomechanical characteristics selected from the group consisting of:

3. The method described in claim 2, wherein the improved biomechanical property is stiffness.

4. The method of any one of claims 1 to 3, wherein the temperature of the PLA-collagen scaffold before sterilization is essentially the same as or higher than the temperature of the PLA-collagen scaffold during sterilization.

5. The method according to any one of claims 1 to 4, wherein the sterilization is carried out at a temperature in the range of -200°C to 40°C.

6. The method according to any one of claims 1 to 5, wherein said filling in step i) is carried out with a collagen gel.

7. The method of claim 6, wherein the concentration of the collagen in the gel is about 0.1% to 2.0% w / w.

8. 8. The method according to any one of claims 1 to 7, wherein the scaffold obtained in step i) comprises 5-25% w / w collagen, as a percentage relative to the total amount of PLA and collagen.

9. The method of any one of claims 1 to 8, wherein the collagen is recombinant collagen, tissue-derived collagen, or a combination thereof.

10. The mesh used in step i) of claim 1, comprising fibres of PLA, i) providing PLA in solid form; ii) subjecting the PLA to a process, thereby obtaining fibers of PLA; and iii) subjecting the obtained fibers to a process, thereby obtaining a mesh of fibers, Step ii) is carried out by spinning, and / or subjecting said mesh of step iii) to a process comprising carding or needle punching to obtain a 3D network. The method according to any one of claims 1 to 9.

11. The method of claim 10, wherein step ii) is carried out by electrospinning or melt spinning.

12. The method of claim 10 or 11, wherein the collagen in the PLA-collagen scaffold is crosslinked.

13. 13. The method according to any one of claims 1 to 12, wherein a medical sterile scaffold is loaded with a drug substance, and sterilization of the scaffold is performed before or after loading the scaffold with the drug substance.

14. The method of claim 13, wherein the drug substance is an anti-inflammatory drug.

15. The method of claim 13, wherein the drug substance is ibuprofen.

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