Method for preparing a three-dimensional scaffold for medical use

Abstract

A method for preparing a sterilized stent for medical use, the method comprising the steps of: i) loading collagen onto a fiber mesh containing poly(lactide) polymer or copolymer (commonly referred to as PLA) fibers to obtain a PLA-collagen stent, ii) drying the PLA-collagen stent obtained from step i), iii) sterilizing the PLA-collagen stent obtained from the drying step ii) to obtain a sterilized stent. The obtained sterilized stent has improved biomechanical properties compared to the non-sterilized stent.

Description

Technical Field

[0001] The present invention relates to a method for preparing a biomaterial for use as a three-dimensional scaffold in a medicament or a cosmetic. The scaffold consists of a poly(lactide) or polylactic acid polymer or copolymer (commonly referred to as PLA) and collagen. The method involves steps enabling the preparation of a scaffold having desired properties regarding: i) a set of biomechanical characteristics, ii) stability and iii) purity. Furthermore, the method does not include steps that can cause significant degradation of the components used (i.e., PLA and collagen), or steps that can cause significant degradation of the obtained PLA-collagen scaffold. Moreover, the sterilization step has unexpectedly proven to confer certain biomechanical properties to the scaffold Background Art

[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 (promoting cell adhesion, growth, and tissue formation). Such biomaterials generally help the body to reconstruct damaged tissues and ultimately minimize associated pain and healing times. Particularly for tissues where load-adaptive properties are essential, the combined static and dynamic biomechanical properties of the scaffold are crucial for the ultimate success of the treatment. Any progress in scaffold development should ensure a high correlation between in vitro conditions and expected in vivo tissue regeneration. The non-toxic biodegradation of the scaffold should gradually transfer stress to the newly growing tissue over an appropriate time period. The synergistic effect of the correct mechanical stimuli depends largely 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 progress not only with age, obesity, or systemic diseases, but also in young and active populations due to physical causes (e.g., injury). If untreated, these defects can progress to osteoarthritis (OA), affecting more than 240 million people worldwide. In full-thickness cartilage defects, natural wound healing often leads to the formation of fibrocartilage, which is both functionally and biomechanically inferior to the original hyaline cartilage, making the tissue more prone to further joint deterioration and osteoarthritis changes. The resulting vicious cycle ultimately leads to total or partial joint replacement. Therefore, a biomaterial solution with the ability to regenerate cartilage is highly desirable for treating early cartilage lesions before the onset of OA

[0004] Biomaterials used clinically include various materials of natural origin and synthetic materials. The advantage of natural materials lies in their natural feasibility for this purpose, although the use of materials of animal origin (xenografts) contains certain risks, such as contamination and unwanted immune responses. This can be avoided by using synthetic materials that do not themselves cause foreign body or allergic reactions. Synthetic materials can have greater biological advantages and biocompatibility. On the other hand, compared with materials of natural origin, synthetic polymers generally lack the desired intrinsic biological cues to promote cell adhesion, proliferation, and tissue restoration. However, it is challenging to evaluate and optimize any biomaterial for clinical use and for the purpose of "precision medicine" solutions. It is currently generally believed that the current level of evaluating the mechanical functions of biomaterials and tissue-engineered constructs is far from sufficient.

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

[0006] The structure, function, and biomechanical behavior of AC are very complex, highly anisotropic, and time- and load-history dependent. Articular cartilage consists of a relatively small number of chondrocytes surrounded by a multi-component matrix, which can be imaged as a complex with 70-85% water and the remaining proteoglycans (proteins with glycosaminoglycans attached in a bottlebrush structure) and collagen. The concentration of proteoglycans and water varies with the depth of the cartilage tissue.

[0007] In the absence of blood supply and lymphatic drainage, articular cartilage is in isolation and almost lacks the wound healing response of other connective tissues. The tissue is highly exposed to biomechanical distortions, resulting in a high incidence of cartilage lesions. Such lesions caused by trauma or due to long-term non-physiological loading often progress to osteoarthritis (OA). OA is the leading cause of musculoskeletal diseases globally, with an incidence of 7-10% in Western populations. The estimated cost of OA for newly diagnosed patients is $6,800 per year, so delaying OA by 10 years results in a savings of $68,000 per patient. The expenditure on OA in the EU is approximately 15-20 billion per year. Although not traditionally applicable to the treatment of OA, cartilage repair has become an increasing focus of attention due to its potential to alter the progression of degenerative diseases and the promise of delaying or eliminating the need for joint replacement.

[0008] In addition to the significant morbidity and the potential for disability, cartilage trauma and degeneration also have a major economic impact. It is estimated that the incidence of cartilage lesions is 23 / 100,000 people per year, and in the European Union, more than 100,000 patients with knee cartilage defects require repair treatment. Due to the aging population and the increasing obesity rate, the prevalence of cartilage lesions is expected to increase rapidly in the coming decades; it is expected that the demand for knee replacement will increase significantly by 2030. On the other hand, due to the high functional requirements and limited treatment options, young patients with symptomatic cartilage lesions represent a challenging group. The aim of articular cartilage repair treatment is to restore and maintain the normal function of the joint with a repair tissue structure that is indistinguishable from native hyaline cartilage. However, the current repair techniques for cartilage lesions are not yet perfect and need further development.

[0009] After surgical repair, the biomechanical properties at the repair site are weakened, and postoperative loading must be reduced. Therefore, it is expected that the lack of mechanical stimulation leads to slow tissue turnover and healing; thus, the recovery time remains long. As mentioned above, biomaterial scaffolds can provide structural support for the healing lesion to enable early weight-bearing and thereby enhance the healing process. A wide variety of three-dimensional scaffolds (including both natural and synthetic ones) have been introduced for cartilage repair.

[0010] A class of scaffolds of particular interest in this context are scaffolds made of poly(lactide) polymers or copolymers (commonly referred to as PLA). Such scaffolds have been the subject of several publications:

[0011] For example, PLA scaffolds have been described in Muhonen et al., published online in Wiley Online Library DOI 10.1002 / jor23099, 2015. Gamma-irradiated PLA scaffolds were immersed in a solution of recombinant human type II collagen to obtain rhCo-PLA scaffolds, and the scaffolds were tested in a porcine study using the matrix-induced autologous chondrocyte implantation (MACI) procedure. The results were compared with those of a MACI-treated group using a commercial membrane and a control group that received no treatment. Compared with the control group, both treatment groups showed improvement, but the differences were not statistically significant. In some rhCo-PLA-treated lesions, the mechanical properties as well as the repair tissue structure were very similar to those of healthy cartilage.

[0012] Gasik et al., Bioengineering and Biotechnology, Volume 6, November 2018, involves sterile PLA scaffolds doped with recombinant human collagen III solution to obtain rhCo-PLA scaffolds. Two types of scaffolds were compared: PLA and rhCo-PLA scaffolds. The results of the biomechanical comparison of the two scaffold types showed how collagen addition and the composition of the medium altered the elastic, viscoelastic, and inelastic properties of the scaffolds. In this study, the analysis of the in vivo cartilage repair scores from the study by Muhonen et al. was also performed using BUGS - Bayesian inference with Gibbs Sampling, i.e., a form of Markov Chain Monte Carlo sampling. The results using the normal or Poisson distribution of the standardized 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 scaffolds and 0.288 for the spontaneous healing control group.

[0013] Salonius et al., J. Cell. Physiol. 2019 involves cell therapy combined with biomaterial scaffolds. The experimental scaffold type used was based on recombinant human collagen - poly(lactide) (rhCo-PLA), which was prepared by immersing sterile PLA scaffolds in a solution of recombinant human collagen II or III. The control scaffold type was a commercial soft membrane produced from porcine-derived type I / III collagen. The results showed that type II collagen promoted the proliferation of mesenchymal stromal cells (MSCs), but the type of collagen used in the rhCo-PLA scaffolds did not affect MSC differentiation during in vitro culture, while chondrogenic differentiation of MSCs led to cell hypertrophy in both rhCo-PLA scaffolds and commercial collagen membranes. However, it was claimed that the limitations of static cell culture might have influenced the results, as hypertrophy is a common phenomenon in MSC in vitro chondrogenesis, while MSCs in in vivo articular cartilage defects did not show upregulation of hypertrophy markers. In joints, cells are under cyclic mechanical loading, which provides a large number of nutrients and differentiation cues to the joint.

[0014] WO 2016 / 042211 (University of Helsinki et al.) involves a three-dimensional material obtained by immersing a sterile polymer felt in a collagen solution. The polymer can be PLA. It has been shown that such scaffolds have improved retention / stiffness compared to scaffolds containing bovine type I collagen or recombinant human type II collagen. Compared to the Compared with the scaffolds, rhCo-PLA showed better performance after 4 months of healing following implantation in the right knee of pigs.

[0015] However, there is clearly a need for cost-effective, safe and reliable scaffolds that have the desired properties and are still safe and sterile. In addition, it would be advantageous if such scaffolds had general utility not only in cartilage repair but also in general medicine or cosmetology, including human and veterinary uses. Summary of the Invention

[0016] The present invention relates to a method for preparing a sterilized scaffold for medical use, the method comprising the following steps:

[0017] i) Loading collagen onto a fiber mesh containing poly(lactide) or copolymer (commonly referred to as PLA) fibers to obtain a PLA-collagen scaffold,

[0018] ii) Drying the PLA-collagen scaffold obtained from step i),

[0019] iii) Sterilizing the PLA-collagen scaffold obtained from the drying step ii) to obtain a sterilized scaffold.

[0020] During the preparation, measures are taken to avoid unnecessary degradation of PLA, collagen and the PLA-collagen scaffold. Therefore, the sterilization step of the PLA-collagen scaffold is carried out without giving any or at least only a slight temperature rise in the scaffold. This can be achieved, for example, by selecting the sterilization method and / or by taking special precautions to avoid temperature rise. In addition, as shown in Example 3 herein, the sterilized scaffold has improved biomechanical properties compared to the non-sterilized scaffold, and thus also has improved biomechanical properties compared to the scaffold prepared under aseptic conditions.

[0021] The fiber mesh used in the method of the present invention can be obtained by the following steps:

[0022] i) Providing PLA in solid form,

[0023] ii) Subjecting the PLA to a process to obtain PLA fibers, and

[0024] iii) Subjecting the obtained fibers to a process to obtain a fiber mesh.

[0025] The fibers of PLA can be produced by known methods. In an embodiment of the present invention, a method suitable for producing PLA fibers is by spinning, such as by melt spinning or electrospinning. Spinning is typically carried out by melting PLA and subjecting the molten PLA to a spinning process, or by dissolving PLA in a suitable solvent and subjecting the PLA solution to a spinning process. The obtained fibers can be subjected to a process such that the fibers are incorporated in a web, and such a web can be further subjected to a process to obtain a 3D structure; in addition, the web can be subjected to a process to ensure that the fibers of the web or its 3D structure are held together.

[0026] Different scaffold structures can be obtained.

[0027] 1. One-dimensional (1D) fiber structure, where collagen is loaded on the fibers (the length and diameter of the fibers can vary).

[0028] 2. Impermeable two-dimensional (2D) matrix: This structure allows bioactive substances (such as cells) to be incorporated in the structure. Typically, cells are cultured in a 2D environment.

[0029] 3. Three-dimensional (3D) nanoporous hydrogel scaffold, where collagen is typically located on top of the hydrogel (i.e., interacting with a 2D matrix having a nanoscale surface) or encapsulated in a 3D structure (cells must degrade the surrounding hydrogel to move or expand).

[0030] 4. Three-dimensional (microporous) scaffold, where due to the high porosity (usually >70%) of the scaffold, bioactive substances can diffuse in three-dimensional space, and depending on the pore size of the scaffold, they can be arranged along one-dimensional scaffold struts or attached to multiple struts and diffuse in three-dimensional space.

[0031] In this context, the description of the scaffold is also intended to cover sterilized scaffolds.

[0032] Generally, the scaffolds of the present invention have a three-dimensional structure. As described above, the three-dimensional structure can also be in the form of a nanoporous hydrogel scaffold, where bioactive substances are on top of the hydrogel or encapsulated within the 3D structure, and the surrounding hydrogel must be degraded to release the bioactive substances, or the three-dimensional structure can be in the form of a microporous scaffold, where due to the high porosity of the scaffold, bioactive substances can diffuse in three-dimensional space.

[0033] The scaffold according to the present invention can be used in this way, or loaded with one or more bioactive substances. The bioactive substance can be one or more cells, or it can be an agent that is active in a biological environment, especially in a mammalian body. The cells to be loaded into the scaffold according to the present invention can be cells intended for repairing diseased or damaged tissues. Other bioactive substances can be active ingredients suitable for relieving pain or treating diseases in specific tissues. It can also be an active ingredient intended for systemic use, but an active ingredient that is easily administered into the implant. The scaffold can be in the form of an implant or a bandage. It can be used in medicine, such as human medicine, as well as veterinary medicine. The scaffold according to the present invention can be used for different medical purposes, especially those related to cartilage repair, such as AC, or those related to osteochondral repair. It can also be used in treatment regimens, such as in AO, etc.

[0034] Native cartilage does not have such porosity. However, for example, the porosity of natural bone containing a large amount of collagen is in the range of 50 - 90%, depending on the type of bone. Tissue engineering scaffolds require a highly porous, interconnected, and open pore structure to ensure tissue ingrowth and the flow and transport of nutrients and metabolic wastes. The porosity of materials can be determined by various methods. For example, micro-computed tomography (microCT) analysis, image analysis (such as scanning electron microscopy or transmission electron microscopy), gas pycnometry, and mercury and liquid extrusion porosimetry can be used. For the overall porosity, microCT analysis is considered the best method because it gives reliable results for the determination of the overall porosity.

[0035] The fiber mesh used in this method 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 containing both the PLA fiber mesh and the collagen component is about 70 to 99%, preferably about 80 to 95%. The porosity of the scaffold before and after sterilization is in the same range.

[0036] Detailed description of the present invention

[0037] As mentioned above, the present invention provides a method for preparing a sterilized scaffold. The method includes the following steps:

[0038] i) Loading collagen onto a fiber mesh containing poly(lactide) or copolymer (commonly referred to as PLA) fibers to obtain a PLA - collagen scaffold,

[0039] ii) Drying the PLA - collagen scaffold obtained from step i),

[0040] iii) Sterilizing the PLA - collagen scaffold obtained from the drying step ii) to obtain the scaffold.

[0041] 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 the PLA-collagen scaffold. Thus, the amount / quantity of monomers present in the starting material PLA in this process does not change significantly, such that the amount / quantity of PLA monomers in the final scaffold is close to that of the starting PLA material. During the pre-step of the method of the present invention, the PLA may be subjected to elevated temperatures in order to obtain a fiber mesh containing PLA fibers, and some degradation of the original PLA material is expected to occur, for example, during the spinning process, to obtain the PLA fiber mesh. Generally, the monomer content in the original PLA material (before fiber formation) is very low, for example, about 0.1%, and the monomer content in the fiber mesh is generally at most about 1%.

[0042] Similarly, by the method of the present invention, only minimal degradation of the collagen occurs during the manufacture of the scaffold. In particular, by avoiding an increase in the temperature of the collagen component during manufacture, especially during the sterilization process, and by ensuring that the sterilization method has no extensive negative impact on the biomechanical properties of the scaffold, the stability of the scaffold during its conversion into a sterilized scaffold is ensured. As mentioned previously, the sterilization process unexpectedly has a positive impact on the biomechanical properties of the scaffold. Thus, sterilization by gamma irradiation at room temperature or lower temperatures results in a scaffold with a more stable structure both in the dry and wet states. The scaffold becomes harder and its biomechanical properties change less, which is advantageous in terms of the approval of the prepared scaffold.

[0043] Regarding stability issues, PLA contains ester bonds that are susceptible to thermal, radiological, and hydrolytic degradation. Collagen can also be degraded by a variety of mechanisms, including enzymatic degradation, radiation-induced degradation, or temperature-dependent degradation.

[0044] The obtained scaffold can be used as it is, or it can be loaded with bioactive substances, such as but not limited to cells, or loaded with active pharmaceutical ingredients and used as an implant. In such cases, the sterilization of the scaffold can be carried out before or after loading with the bioactive substance (such as but not limited to active pharmaceutical ingredients or cells).

[0045] The method of the present invention provides a scaffold having one or more of the following characteristics:

[0046] · When used in an implant, it should be biocompatible to support cell viability and biodegradable such that new tissue replaces the scaffold over time and the degradation products are harmless locally or systemically.

[0047] · It should have a highly porous 3D structure with interconnected pores to enable cell migration within the scaffold such that they can proliferate and form new tissue.

[0048] · It should have an appropriate porosity to allow cell seeding into the scaffold before use, such that the cells remain within the scaffold so that they can proliferate and form new tissue.

[0049] · It should have an appropriate biomechanical strength for applications, e.g., for mechanical stimulation to work well in an AC environment to enable new tissue formation.

[0050] · It should not highly mimic the native tissue it is intended to replace, but it should have suitable compression and decompression functions for applications, as well as fluid absorption and interstitial fluid pressurization during loading and compression.

[0051] · It should have an appropriate degradation time for applications to be able to assist tissue ingrowth and subsequent new tissue formation.

[0052] The polymer for use in the method of the present invention

[0053] Prepare a PLA - collagen scaffold according to the present invention. In this context, PLA is intended to include poly(lactide) in stereoisomeric forms, including poly(lactide) based on L - lactide, D - lactide, and polymers based on both L - lactide and D - lactide. The ratio between the L - form and D - form content can vary. Generally, the PLA used in this method contains both L - form and D - form lactide. In the case of poly(lactide), the presence of the D - form in the polymer generally affects the time taken for the polymer to degrade. The content of the D - form shortens the degradation time. In this way, PLA with a desired degradation time can be designed by varying the content of the L - form and D - form in the polymer. However, the content of the D - form also affects the biomechanical properties of the obtained scaffold. Thus, the higher the content of the D - form in the scaffold, the weaker the biomechanical properties. Therefore, it is important to select an appropriate balance between the L - form and D - form in the poly(lactide) used in the method of the present invention. Generally, the content of the D - form is usually about 1 to about 50% w / w, such as 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 poly(lactide) contains 96% of the L - form and 4% of the D - form; such lactide is designated as 96 / 4 poly(L / D) lactide.

[0054] In this context, the term PLA also includes poly(lactide) or copolymers, such as those formed between lactide and glycolide, poly(lactide - co - glycolide) (PLGA). The content of lactide and glycolide can vary.

[0055] The raw PLA material for obtaining the PLA fiber web used in the method of the present invention has an inherent viscosity - measured 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 wt% to about 1 wt%, such as about 0.1 wt% to at most 1 wt%.

[0056] More details regarding biodegradable PLA polymers suitable for use in the method of the present invention are given below.

[0057] Suitable polymers are biodegradable polymers. They can be natural or synthetic polymers. Generally, synthetic bioabsorbable polymers have been widely studied as tissue engineering scaffolds. Their controllable chemical properties and characteristics, as well as their ease of replication, have gone beyond their use as scaffold materials. Depending on the susceptibility of their functional groups to hydrolysis, synthetic bioabsorbable polymers can be divided into various subgroups, such as esters, orthoesters, anhydrides, carbonates, and amides. In particular, polyesters have been used in many clinical applications because they are easily degraded by hydrolysis of the ester bond. Additionally, their degradation products are in some cases absorbed through metabolic pathways, and they have the potential to change their degradation rate by adjusting their structure.

[0058] Some of the most widely and earliest studied synthetic bioabsorbable polymers for tissue engineering are polyesters. Poly(α-esters) are unique in their great diversity and synthetic versatility. Within the poly(α-ester) category, poly(α-hydroxy acids), including polyglycolide (PGA) and stereoisomeric forms of polylactide (PLA), are the most widely studied polymers.

[0059] In the method of the present invention, PLA is used. PLA is a thermoplastic biodegradable polymer produced by the polycondensation of lactic acid fermented from sugars derived from carbohydrate sources (such as corn, sugarcane, and cassava), or by ring-opening polymerization from lactide (the cyclic dimer of lactic acid). Due to its chiral carbon atom, lactic acid exists in two enantiomeric forms, called L-lactic acid (S) (which occurs in the metabolism of all animals and microorganisms) and D-lactic acid (R). By polycondensation, usually only low molecular weight PLA can be obtained. High molecular weight PLA can be obtained by ring-opening polymerization, in which lactic acid is polycondensed and then depolymerized into the dehydrated cyclic dimer, lactide. Optically active lactide can exist as D-lactide, L-lactide, or meso-lactide (D,L-lactide). In addition to the three diastereomeric structures, a racemic mixture of D-lactide and L-lactide, racemic lactide, also exists. The structure and composition of the polymer chain and especially the ratio of L- to D-isomers of lactic acid affect the processing, crystallization, and degradation behavior of PLA. By copolymerizing L-lactide with meso-L,D- or racemic lactide, high molecular weight amorphous or semi-crystalline polymers with a melting point range of 130 to 185 °C can be obtained. Poly(L-lactide) (PLLA), i.e., a homopolymer containing only L-lactide, is semi-crystalline and has the highest melting point, while PLA copolymers with a higher D-isomer content exhibit lower melting points and significantly lower crystallization behavior, becoming amorphous when the D-content is higher than 12 - 15%.

[0060] PLA is an aliphatic polyester and is thus susceptible to hydrolysis degradation because of the presence of ester groups in its structure. The hydrolysis degradation behavior, rate, and mechanism can be controlled by changing the molecular and higher-order structures as well as medium factors such as the temperature, pH, and catalytic substances (such as bases and enzymes) of PLA. The in vivo hydrolysis degradation rate is comparable to that in vitro, so the in vivo degradation can be predicted to a certain extent from the in vitro degradation behavior and rate. PLA does not require the presence of enzymes to catalyze hydrolysis. Lactic acid occurs in the metabolic processes of living organisms, so the degradation products of PLA are non-toxic. The hydrolysis of aliphatic polyesters begins with water entering the matrix, followed by the hydrolytic cleavage of the ester bond. The initial crystallinity affects the hydrolysis degradation rate because the amorphous part has a higher water absorption rate, and the crystalline segments reduce the water penetration in the matrix. In addition, autocatalysis of PLA samples has also been reported. Autocatalysis is due to an increase in the number of compounds with carboxylic acid end groups in the center of the sample when low molar mass compounds cannot penetrate the outer shell (where the degradation products dissolve in the surrounding solution).

[0061] Due to its thermoplastic nature, PLA can be processed into various forms. Melt processing is the most widely used method for PLA. Additionally, injection molding and extrusion are widely used methods for manufacturing PLA films and fibers for different non-woven fabrics or textiles. Moreover, electrospinning of PLA is also used in medical applications to produce thin fibers, which can be used as medical tissue scaffolds, wound dressings, drug carriers, protective fabrics, and nanocomposites. The extensive medical applications of PLA include orthopedic screws, tissue-engineered scaffolds, sutures, protein encapsulation and delivery, microspheres, and drug delivery systems. PLLA is a slow-degrading polymer with good tensile strength, low elongation, and high modulus (complete absorption in the body takes more than 2 to 5 years). This is why PLLA is considered ideal for load-bearing applications such as orthopedic fixation devices. On the other hand, PDLLA degrades faster and loses strength within 1 - 2 months and experiences mass loss within 12 - 16 months when hydrolyzed. Its tensile strength is also lower compared to PLLA. Given this, PDLLA is preferably used as a drug delivery carrier and low-strength scaffold material for tissue engineering. PLLA (semicrystalline), PLDLA (amorphous), P(L / DL)LA 70 / 30 (amorphous), and P(L / D)LA 96 / 4 (semicrystalline) are the most commonly used PLA polymers in the medical industry.

[0062] Polylactide-based copolymers

[0063] PLGA is the most studied biodegradable polymer in biomedical applications. Since PLA and PGA have significantly different properties, different copolymer compositions allow PLGA to be optimized for different applications. PLGA with a lactide composition of 25 - 75% forms an amorphous polymer, which is very unstable in terms of hydrolysis compared to the more stable homopolymers. Many different processing techniques have been used for PLGA scaffold fabrication, such as gas foaming, microsphere sintering, porogen leaching, electrospinning, and polymer printing. Since PLGA degrades faster compared to other polyesters, PLGA is particularly used as sutures and drug delivery devices. PLGA is also made into tissue-engineered scaffolds because it shows good cell adhesion and proliferation properties.

[0064] As mentioned above, in this context, the term PLA covers all polylactide polymers (also known as polylactic acid polymers) and copolymers having lactide and glycolide. In a specific embodiment, PLA is a polylactide polymer.

[0065] Regarding stability, PLA contains ester bonds and can be degraded into lower molecular weight PLA, or into monomers, dimers, etc. In the present method, measures are taken to avoid extensive degradation. The degradation rate depends on temperature and pH. In the method of the present invention, after melt spinning, at least some or all of the process steps are carried out mostly at room temperature, i.e., mostly at 25 - 30 °C, for example mostly at 25 °C, and several process steps (such as some steps in drying and sterilization) are carried out at temperatures significantly lower than room temperature, for example at 0 °C or even lower, such as not greater than -10, -20 or -25 °C.

[0066] By varying the intrinsic viscosity or the monomer amount, the stability of the PLA component of the PLA - collagen scaffold can be ensured after processing (see, for example, Example 1 herein). The overall stability of the PLA - collagen scaffold can be ensured after all processing steps, for example, as shown in Examples 2 and 3 herein by biomechanical characterization.

[0067] Definitions

[0068] As used herein, "three - dimensional material" or "three - dimensional structure" refers to any material having height, width, and depth. An example of a three - dimensional structure is a scaffold. The three - dimensional materials of the present invention are preferably implantable, biodegradable, and biocompatible.

[0069] As defined herein, a "biodegradable material" is a material that does not require retrieval or further manipulation after introduction into the body because it degrades into soluble and non - toxic by - products.

[0070] As defined herein, an "implantable material" is any material of any shape or size that is suitable for implantation into a subject.

[0071] As defined herein, a "biocompatible material" is a material that is harmless or non - toxic to living tissue.

[0072] As defined herein, "loading", such as loading collagen into a fiber mesh, is intended to represent a process by which collagen is added to the fibers or brought into contact with the fiber mesh such that the collagen can be present on top of the fiber mesh, or incorporated into the fiber mesh or both, or the fiber mesh is impregnated with collagen.

[0073] The term "biomechanical strength" as used herein is intended to refer to the ability of a scaffold applied to biological tissue to maintain its function without breaking into pieces, and the ability of the scaffold to withstand normal handling of the scaffold during manufacture, storage, and application.

[0074] The term "mechanical strength" as used herein is intended to refer to the ability of a scaffold to withstand normal handling of the scaffold during manufacture, storage, and use. In this context, "mechanical strength" is sometimes used synonymously with "biomechanical strength" and with "biomechanical function".

[0075] In this context, the 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 the scaffold prepared according to the present invention compared to a scaffold prepared in the same manner but in which the last sterilization step is omitted. The desired improvement is when the increase (or decrease) is 1% or greater, 2% or greater, 3% or greater, 4% or greater, 5% or greater, 6% or greater, 7% or greater, 8% or greater, 9% or greater, or 10% or greater; or when the 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%. The data used is based on the average of two or more measurements.

[0076] In this context, the relevant biomechanical characteristics or biomechanical parameters include: storage modulus, storage creep modulus (i.e., the storage modulus under creep conditions), dynamic storage modulus, memory value, fluid mobility, apparent permeability, creep permeability (i.e., the apparent permeability under creep conditions), dynamic modulus (i.e., stress / strain ratio), and stiffness (i.e., the stress / strain ratio of the material). The experimental section herein further defines the parameters.

[0077] In this context, the term "creep test" is explained as follows: The nature of the creep test is pseudo-static (the change of strain with time under a continuously applied stress), and it is often used to evaluate the viscoelastic properties of materials and approximate them with some models.

[0078] As used herein, the term "fiber" refers to a material made of fibers. Fibers having only one size or different sizes in diameter can be used to prepare the mesh used in the method of the present invention. These polymeric fibers can be selected from PLA fibers having the following diameters: 5 to 100 μm, such as 5 to 75 μm, 5 to 50 μm, 5 to 40 μm, 5 to 35 μm, 10 to 75 μm, 10 to 50 μm, 10 to 40 μm, 10 to 35 μm, 15 to 75 μm, 15 to 50 μm, 15 to 40 μm, or 15 to 35 μm. The diameter is the average diameter of the fibers in the structure. The cross-section of the fiber is not limited to a circular cross-section, but can also be any other shape, such as oval, star-shaped, rectangular, or triangular.

[0079] In this context, "porosity" is calculated as follows: Porosity = pore volume / sample volume x 100%. In addition, porosity is considered to be evaluable by micro-computed tomography (microCT), where the overall 3D structure of the structure can be analyzed.

[0080] In this context, a "bioactive substance" is a substance, compound and / or active agent having biological or pharmacological activity, i.e., an effect on a living organism, tissue and / or cell, such as but not limited to the beneficial or adverse effect of a drug on a living substance. When a drug is a complex chemical mixture, this activity is exerted by the active ingredient or pharmacophore of the substance, but may be modified by other ingredients. Typical examples are antibiotics, enzymes and vitamins, grafts and cells.

[0081] In this context, the term "about" is intended to denote a range corresponding to -10% to +10% of the specified value.

[0082] Methods for obtaining fibers and fiber webs

[0083] Fibers of PLA polymers can be obtained by various methods well known to those skilled in the art. They can be obtained by melt processing or electrospinning. As can be seen from the examples herein, melt spinning has proven suitable for preparing PLA fibers for use in the methods of the present invention.

[0084] The melt spinning process involves melting the polymer or heating the polymer to a soft form. Accordingly, a suitable temperature is selected for the spinning process depending on the choice of the particular polymer. Generally, when using polylactide, the temperature is in the range of about 60 to about 300 °C, for example in the range of about 70 to about 250 °C.

[0085] Polylactide can be amorphous or semi-crystalline, depending on the ratio between L- and D-lactide monomers. Poly(glycolide) 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 the semi-crystalline poly(L,D-lactide) has an approximate melting unit of 135 - 170 °C. Amorphous polymers do not have a melting point. Poly(glycolide) is produced by ring-opening polymerization of glycolide. It has a crystallinity of about 45 - 55%. It has a high melting point (∼225 °C) and a glass transition temperature of ∼35 °C. Poly(glycolide) degrades relatively rapidly to acidic products.

[0086] Generally, the melt spinning process of PLA is carried out at a temperature in the range of about 60 °C to 300 °C.

[0087] The PLA raw material is dried before spinning and a protective gas is used to prevent degradation during the spinning process.

[0088] The melt spinning process results in PLA fibers that are suitable for use in the PLA fiber meshes of the present invention. During the spinning process, some changes may occur in the PLA material, but these changes should not significantly affect the suitability of using the resulting fiber meshes to obtain PLA-collagen scaffolds. The resulting fibers are semi-crystalline. Generally, the resulting fibers have a monomer content that is close to the initial monomer amount of PLA, and / or a decrease in the inherent viscosity of less than 30%, 25%, or 20% compared to the raw material. This is especially applicable in cases where the inherent viscosity of the PLA raw material is up to 2.5 dl / g. However, if the inherent viscosity of the PLA raw material is 2.5 to about 5 dl / g, the inherent viscosity of the resulting fibers may decrease by 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, if the range of the inherent viscosity of the fibers produced from the spinning process is 1.5 to about 5 dl / g, such as 1.5 to 4 dl / g or about 1.5 to about 3 dl / g, a suitable scaffold is obtained. The inherent viscosity is determined as described herein at room temperature.

[0089] When the fibers have been obtained, they are converted into a mesh. As described above, the mesh can be carded into a 3D network. Carding is a mechanical process that loosens and intermingles the fibers to produce a continuous randomly oriented network, i.e., a carded web or fiber network. Carding breaks the fiber lock-ups and unorganized bundles and then aligns the individual fibers so that almost all of them are separated from each other. To obtain the desired porosity of the network thus obtained, the network can be subjected to needling, which is a process that uses needles with grooves along the needle axis to grasp the upper layer of fibers as the needles enter the fiber mesh / mat and wind them together with the inner layer of fibers. Needling produces a wound and compressed 3D network from the carding and improves the mechanical properties while still leaving the structure highly porous.

[0090] The resulting fiber mesh or fiber 3D network 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 to 99%. As defined herein, porosity, i.e., void fraction, is a measure of the void (i.e., "empty") space in a material and is the fraction of the void volume divided by the total volume. The obtained mesh or 3D network has a porous structure, particularly a network of pores throughout the material.

[0091] The fiber mesh or 3D network can be cut into the desired form and / or size.

[0092] The thickness of the obtained 3D network is generally 0.1 to 50 mm.

[0093] Load collagen onto the fiber mesh

[0094] Before using a collagen-loaded fibrous mesh or 3D network, the fibrous mesh or 3D network can be subjected to a washing procedure. The washing procedure is typically carried out in an aqueous medium, in an alcohol-aqueous medium, or in an alcohol such as ethanol. The washing procedure can be carried out several times, and then the mesh or 3D network is dried. Drying can be carried out by lyophilization or drying at room temperature (e.g., in a laminar flow hood). The fibrous mesh or 3D network can also be cut into pieces of a size suitable for the intended use.

[0095] The obtained fibrous mesh or 3D network is loaded with collagen. The loading can be carried out to ensure that the main loading is on the outer surface of the mesh or 3D network, or the loading can be carried out to load the collagen into the voids of the mesh or 3D network. Collagen can also adhere to the fibers of the mesh or 3D network. It is possible that collagen is present in the mesh or 3D network by combination, for example, collagen can be present on the surface by adhesion, collagen can be present in the voids by capillary forces and / or by adhesion on the surface. It is contemplated that the mechanism by which collagen is incorporated into the scaffold is not of primary concern, but in this context, collagen is considered to be present in the voids and on the surface of the mesh or 3D network.

[0096] Typically, the loading of collagen involves dissolving or dispersing the collagen in a suitable medium. Usually, the medium is an aqueous medium. The pH and / or viscosity of the aqueous medium containing collagen can be adjusted.

[0097] The concentration of collagen in the suitable medium is from about 0.1% to about 5% w / w, such as from about 0.1% to 4% w / w, 0.1% to about 3% w / w, about 0.1% to about 2% w / w.

[0098] Then the mesh or 3D network can be immersed in the collagen-containing medium. Alternatively, the mesh or 3D network can be immersed in a solvent (which does not dissolve the scaffold), and the solution / dispersion of collagen is added to the immersed mesh or 3D network. Collagen can also be loaded into the mesh or 3D network by injecting or spraying the collagen-containing medium into the mesh or 3D network.

[0099] The pH of the collagen aqueous medium should be maintained below 8 during processing. The process is preferably carried out at room temperature (RT), but can also be carried out at an elevated temperature not exceeding 45°C. The weight ratio of PLA:collagen is about 95 wt% to about 60 wt% of PLA and 40 wt% to about 5 wt% of collagen, and preferably about 90 wt% to 75% of PLA and 25 wt% to 10 wt% of collagen. The weight ratio of PLA to collagen (PLA / collagen) is thus 60 / 40 to 95 / 5, such as 70 / 30 to 90 / 10, 75 / 25 to 80 / 20, or 80 / 20 to 90 / 10. Before possible loading with cells, drugs, etc., the porosity of the PLA-collagen scaffold is about 70% to 99%, and preferably about 80% to 95%.

[0100] As defined herein, recombinant human collagen refers to human collagen polypeptides produced by using recombinant techniques, such as using appropriate polynucleotides, expression vectors, and host cells. Recombinant techniques are well known to those skilled in the art, and there are, for example, several commercial recombinant human collagens on the market. The use of recombinant human collagen reduces the risk of transmitting pathogens of known and unknown animal origin and adverse immune responses. In addition, unlike other cartilage regeneration materials of natural origin, recombinant human collagen is not affected by batch-to-batch variability. Thus, recombinant human collagen can be produced in high quantities and uniform quality at the grades required by good manufacturing practice (GMP).

[0101] Collagen is the most abundant protein in the extracellular matrix (ECM) and is a major component of skin and musculoskeletal tissues. To date, at least 28 different types of collagen have been identified. All collagens have a triple-helix structure in which three independent chains, each in a left-handed polyproline II-helix, coil together to form a right-handed, ropelike supercoiled triple helix. In this structure, collagen shows a characteristic repeating sequence, glycine-X-Y, where glycine is a small enough amino acid to pack into the center of the triple-helix structure. The X and Y positions can be any amino acid, but X is usually proline and Y is usually hydroxyproline. Collagens can be divided into different groups based on their structure and supramolecular organization. These groups are fibril-forming collagens, fibril-associated collagens with interrupted triple helices (FACIT), network-forming collagens, anchoring fibrils, transmembrane collagens, basement membrane collagens, and others, each with unique functions. The most abundant group of collagens is the fibril-forming collagens, which account for approximately 90% of the total collagen. Type I collagen is the most abundant and well-studied collagen and forms more than 90% of the organic mass of bone and is the major collagen in tendon, skin, ligament, cornea, and many interstitial connective tissues, with the exception of only a few tissues such as hyaline cartilage, brain, and vitreous body. On the other hand, type II collagen is characteristic and a major component of hyaline cartilage, but it is also present in the vitreous body, corneal epithelium, notochord, nucleus pulposus of the intervertebral disc, and embryonic epithelial-mesenchymal transition.

[0102] Collagen has high mechanical strength, good biocompatibility, low antigenicity, and crosslinking ability, which enables customization of the mechanical, degradation, and water absorption properties of collagen. When fabricated into highly porous scaffolds, crosslinking of collagen is necessary for making scaffolds with sufficient mechanical properties and degradation rates. Collagen has been widely studied for various medical applications and a wide range of tissue engineering applications, such as bone, cartilage, and skin tissue engineering. Modifying collagen with other biodegradable polymers or combining collagen with other biodegradable polymers enhances the potential of collagen as a biomaterial. The most predominant collagen that can be easily prepared in pure form in commercial quantities is type I collagen. It is the most widely used collagen in tissue engineering applications. Different from biodegradable synthetic polymers, collagen scaffolds provide an alternative way to provide biological information to growing constructs. When using collagen scaffolds, extensive cell adhesion and other signals can be achieved, which will improve the quality of tissue engineering products. Generally, freeze-drying or stereolithography methods are used to fabricate porous collagen-based scaffolds, and carbodiimide is commonly applied for crosslinking. Porous collagen scaffolds are usually combined with other components, such as bioceramics or synthetic biodegradable polymers. Collagen has been used in a variety of commercial medical products, such as bioprosthetic heart valves or as wound dressings. The use of collagen can be divided into two different categories: tissue-based devices, where natural stabilized tissues are used as devices (bioprosthetic heart valves), and purified collagen, where collagen is solubilized through an enzymatic digestion step and redissolved into various products (wound dressings). Recombinant collagen is evolving because they provide a way to produce high-purity and disease-free collagen, which can produce all types of collagen, even those with very low abundance in natural tissues.

[0103] The collagen applicable to the methods of the present invention is collagen from any source, including animal-derived, human, recombinant, or synthetic collagen. The collagen can be type I, II, III, IV, V, VI, IX, or XI collagen. Any combination of these collagen types can also be utilized. More specifically, the collagen is type I, II, or III collagen. The collagen can also be a combination of at least type I, II, and III collagen, at least type I and III collagen, at least type I and II collagen, or at least type II and III collagen. Any collagen can also be used in combination with animal-derived, recombinant, or synthetic collagen, or the collagen can be used alone or in combination with other collagens.

[0104] As used herein, "recombinant human collagen material" refers to any material (e.g., any solution or gel) containing recombinant human collagen.

[0105] The collagen used in the method of the present invention can be porous. For example, freeze-drying renders the collagen porous and elastic and is thus well-suited for its purpose, such as to support chondrocyte proliferation and cartilage matrix production. Collagen, such as a freeze-dried collagen network, is an excellent microenvironment for cell attachment. In an embodiment of the present invention, the collagen is freeze-dried. Collagen (e.g., in the form of a collagen solution) can be freeze-dried as such. The pore size of the collagen structure varies between 20 - 250 μm and can be selected from 20 - 250 μm, 50 - 250 μm, 30 - 200 μm, 40 - 200 μm, 50 - 200 μm or 60 - 200 μm. Additionally, prior to freeze-drying, the collagen can be converted into a gel, i.e., one or more collagens can be in the form of a freeze-dried gel.

[0106] Collagen is used in scaffolds, which can be used in tissue repair and regeneration, whether in sponges, sheets or gels. It is believed that collagen scaffolds have the right properties to enable tissue regeneration, such as pore structure, permeability and hydrophilicity. 3D collagen scaffolds support cell adhesion, robust cell expansion, proliferation and differentiation.

[0107] The thus obtained PLA - collagen scaffold can then be dried. Drying can involve freeze-drying, optionally preceded by freezing the wet scaffold at -20 to -50 °C for 10 h. Freezing depends on the materials used and the sample size and generally takes 10 hours. The freezing temperature affects the size of the pores formed in the structure and also depends on the materials used. The freezing temperature generally varies between -10 °C and -80 °C and preferably between -20 °C and -50 °C. The freeze-drying process is typically carried out at the same temperature range as the freezing process or at a lower temperature. The freeze-drying time depends on the materials used and the sample size and is typically between 24 hours and 48 hours.

[0108] After loading with one or more collagen meshes or 3D networks and optionally drying, a PLA - collagen scaffold is obtained, which is preferably subjected to crosslinking of the collagen in the PLA - collagen scaffold. Crosslinking is carried out to increase the biomechanical strength of the final scaffold and to render the collagen component more stable in vivo.

[0109] Suitable crosslinking methods are well-known to those skilled in the art and include, but are not limited to, the use of chemical crosslinking agents such as 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC), glutaraldehyde, genipin and also ultraviolet light, or combinations thereof.

[0110] A suitable combination of crosslinking agents is EDC, and preferably EDC is used together with N-hydroxysuccinimide (NHS). Crosslinking is usually carried out in an alcoholic medium, such as in an ethanol-water solvent. The content of ethanol is usually 50 wt% to 99 wt%, such as 60 wt% to 99 wt%, 70 wt% to 99 wt%, 80 wt% to 99 wt% or 90 wt% to 99 wt%.

[0111] Crosslinking is usually carried out by using EDC and NHS. The concentration range of EDC can be 5 to 30 mM, such as 5 to 25 mM, 10 to 20 mM, such as 14 mM EDC. The concentration range of NHS can be 1 to 15 mM, such as 1 to 10 mM, 5 to 10 mM or about 6 mM NHS.

[0112] The crosslinked PLA-collagen scaffold is usually extensively washed with an aqueous medium, including water, an alcoholic aqueous medium or alcohol, to avoid crosslinking residues.

[0113] Then the crosslinked PLA-collagen scaffold is dried. Drying can include freeze-drying. Optionally, the previous step is to freeze the wet scaffold at -20 to -50 °C for 10 hours. The details of freeze-drying and freezing are the same as those mentioned above. Temperature can affect the size of pores in the final product; thus, the lower the temperature, the smaller the pore size.

[0114] Sterilization of the dried PLA-collagen scaffold

[0115] In addition to being safely used and working properly in, for example, an AC environment, the scaffold should be sterile before implantation. One of the most reliable sterilization methods for implantable medical devices is gamma irradiation. Gamma irradiation is very effective and there are no residual chemicals that can cause cytotoxicity. However, it is known that gamma irradiation affects the properties of biodegradable polymers, such as PLA, and affects collagen. In particular, at overly high levels of irradiation, collagen can suffer a loss of mechanical integrity. Collagen is temperature-sensitive and high temperature increases should be avoided during processing as well as during the sterilization process.

[0116] After drying the PLA-collagen scaffold, it is sterilized to form a sterile PLA-collagen scaffold. Any suitable method that does not cause unnecessary degradation of PLA, collagen or the PLA-collagen scaffold can be used. A suitable method is to subject the PLA-collagen scaffold to gamma irradiation, but as mentioned previously, any temperature increase should be avoided to prevent any degradation of the collagen component. Irradiation is a process involving energy transfer, so this process should be carried out under temperature control conditions for sensitive materials, such as collagen. A method for maintaining low temperature is to cool the PLA-collagen scaffold during the irradiation process. Before sterilization, the scaffold is usually cooled to about -200 °C to about 25 °C. By this method, the temperature during sterilization is kept below 40 degrees.

[0117] The temperature of the PLA - collagen scaffold before sterilization is substantially the same as or higher than the temperature of the PLA - collagen scaffold during sterilization.

[0118] Sterilization is generally carried out at a temperature in the range of - 200°C to 40°C, for example up to 30°C or up to 25°C, or at a temperature in the range of - 100°C to 25°C, for example at a temperature of - 70°C, - 40°C, 0°C, 10°C, 20°C or 25°C.

[0119] As can be seen from the examples herein, the doses used - when sterilized by gamma irradiation - are generally in the following ranges: 10 kGy to about 27 kGy, for example 15 to 26 kGy, 16 to 25 kGy, for example 18 kGy, 19 kGy, 20 kGy, 21 kGy, 22 kGy, 23 kGy, 24 kGy, 25 kGy, 26 kGy or 27 kGy.

[0120] Generally, the PLA - collagen scaffold is packaged in a suitable package before sterilization to ensure the sterility of the scaffold after the sterilization process.

[0121] It should be mentioned that sterilization can also occur in an optional step of the method of the present invention. Thus, the fibrous web (PLA web) can be subjected to sterilization and then loaded with collagen, and the collagen itself can be sterilized and then loaded into the PLA fibrous web or 3D network.

[0122] The main advantage of using the sterilization process is that the resulting sterile scaffolds have very uniform biomechanical properties. As shown in Example 3, scaffolds manufactured aseptically but not sterilized have much more variable biomechanical properties. From a regulatory perspective, the main advantage is the ability to produce scaffolds with no significant variation in characteristics from batch to batch. The reason is that biomechanical properties give an indication of how the scaffold will behave in the body, and what is desired is a predictable in - vivo behavior, i.e., no or only minor variations have a minimal impact on in - vivo characteristics, while larger variations can have an impact on in - vivo characteristics and will require investigation of whether such impacts negatively affect the function of the scaffold (e.g., stability of the scaffold, ability to remain at the administration site, negative impact on the release of cells and bioactive agents, unwanted side effects, etc.). Therefore, strict control to obtain the same or almost the same biomechanical properties is desired - also in relation to regulatory approval. Generally, a variation in characteristics of ±10% is usually acceptable.

[0123] As can be seen from the examples herein, the scaffolds prepared according to the present invention have improved biomechanical properties compared to non - sterilized scaffolds.

[0124] Improved biomechanical properties can be represented as a change in one or more biomechanical parameters or biomechanical characteristics, such as the storage modulus, the storage creep modulus (i.e., the storage modulus under creep conditions), the dynamic storage modulus, the memory value, fluid mobility, apparent permeability, creep permeability (i.e., the apparent permeability under creep conditions), the dynamic modulus (i.e., stress / strain ratio), and stiffness (i.e., the stress / strain ratio of the material). The desired change can be i) an increase, ii) a decrease, or iii) no change.

[0125] In one aspect, one or more biomechanical parameters or biomechanical characteristics are selected from the storage creep modulus when tested under wet conditions and the dynamic modulus when tested under dry conditions. The desired change compared to a non-sterilized scaffold is an increase. Generally, the increase is 1% or greater, 2% or greater, 3% or greater, 4% or greater, 5% or greater, 6% or greater, 7% or greater, 8% or greater, 9% or greater, or 10% or greater.

[0126] Some biomechanical parameters can decrease or remain unchanged. Such parameters include biomechanical parameters or biomechanical characteristics selected from creep permeability and dynamic modulus, both determined under wet conditions. The desired decrease is 1% or greater, 2% or greater, 3% or greater, 4% or greater, 5% or greater, 6% or greater, 7% or greater, 8% or greater, 9% or greater, or 10% or greater, 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%.

[0127] In particular, the improved biomechanical property is improved stiffness.

[0128] All details regarding each step mentioned in the main aspects of the present invention apply, mutatis mutandis, to all other aspects of the present invention and are thus not repeated in the following paragraphs.

[0129] Other aspects of the invention

[0130] The present invention also relates to a method for preparing a sterilized PLA-collagen scaffold, the method comprising:

[0131] a) providing PLA in solid form

[0132] b) subjecting the PLA to a process to obtain PLA fibers, and

[0133] c) subjecting the obtained fibers to a process to obtain a fiber web,

[0134] d) optionally subjecting the fiber web to carding and / or needling to obtain a 3D network of PLA fibers

[0135] e) Optionally, subject the fiber web or 3D fiber network to one or more washing procedures,

[0136] f) Provide collagen in the form of a solution or a gel,

[0137] g) Immerse the fiber web or 3D fiber network obtained after step c) or (if included) step d) or e) in the collagen solution or collagen gel to obtain a PLA - collagen scaffold,

[0138] h) Optionally, dry the PLA - collagen scaffold, and

[0139] i) Sterilize the PLA - collagen scaffold obtained in step g) or (if included) step h).

[0140] The present invention also relates to a method for preparing a sterilized PLA - collagen scaffold, the method comprising:

[0141] a) Provide PLA in solid form,

[0142] b) Subject the PLA to a process to obtain PLA fibers,

[0143] c) Subject the obtained fibers to a process to obtain a fiber web,

[0144] d) Subject the fiber web to carding and / or needling to obtain a 3D network of PLA fibers,

[0145] e) Optionally, subject the fiber web or 3D fiber network to one or more washing procedures,

[0146] f) Provide collagen in the form of a solution or a gel,

[0147] g) Immerse the fiber web or 3D fiber network obtained after step d) or (if included) step e) in the collagen solution or collagen gel to obtain a PLA - collagen scaffold,

[0148] h) Optionally, dry the PLA - collagen scaffold, and

[0149] i) Sterilize the PLA - collagen scaffold obtained in step g) or (if included) step h).

[0150] The present invention also relates to a method for preparing a sterilized PLA - collagen scaffold, the method comprising:

[0151] a) Provide PLA in solid form,

[0152] b) Subject the PLA to a process to obtain PLA fibers,

[0153] c) Subject the obtained fibers to a process to obtain a fiber web,

[0154] d) subject the fiber web to carding and / or needling to obtain a 3D network of PLA fibers,

[0155] e) subject the 3D fiber network to one or more washing procedures,

[0156] f) provide collagen in the form of a solution or a gel,

[0157] g) immerse the 3D fiber network obtained after step e) in the collagen solution or collagen gel to obtain a PLA-collagen scaffold,

[0158] h) optionally, dry the PLA-collagen scaffold, and

[0159] i) sterilize the PLA-collagen scaffold obtained in step g) or (if included) step h).

[0160] The present invention also relates to a method for preparing a sterilized PLA-collagen scaffold, the method comprising:

[0161] a) provide PLA in solid form,

[0162] b) subject the PLA to a process to obtain PLA fibers,

[0163] c) subject the obtained fibers to a process to obtain a fiber web,

[0164] d) subject the fiber web to carding and / or needling to obtain a 3D network of PLA fibers,

[0165] e) subject the 3D fiber network to one or more washing procedures,

[0166] f) provide collagen in the form of a solution or a gel,

[0167] g) immerse the 3D fiber network obtained after step e) in the collagen solution or collagen gel to obtain a PLA-collagen scaffold,

[0168] h) dry the PLA-collagen scaffold, and

[0169] i) sterilize the PLA-collagen scaffold obtained in step h).

[0170] Use of the scaffold prepared by the method of the present invention

[0171] The scaffold obtained according to the present invention is generally used for surgical care of humans and veterinarians. It can also be used in other health and medical care and cosmetics for humans and animals (i.e., companion animals).

[0172] The scaffolds obtained according to the present invention can be used for treating lesions in joint surfaces, especially in weight-bearing joints, such as the knee joint. These lesions present symptoms such as pain and locking of the affected joint and require surgical intervention. The etiologies of these lesions are different, but cartilage damage can be the result of trauma, degenerative joint diseases (such as osteoarthritis (OA)), and developmental disorders (such as osteochondritis dissecans).

[0173] The scaffold can be used as such, or it can be loaded or combined with one or more of the following: tissue-specific cells, such as chondrocytes, somatic stem cells or embryonic stem cells, such as bone marrow mesenchymal stromal cells, cellular components, such as growth factors or cytokines, blood components and fractions, such as platelets or platelet-rich plasma, or active pharmaceutical ingredients, such as anti-inflammatory drugs, such as ibuprofen. For the above-mentioned cases, it is crucial to use the scaffold of the present invention as a lesion filler and, of secondary importance, to incorporate potential additives.

[0174] The scaffolds obtained according to the present invention can also be used in cosmetic surgery, for example, as a dermal filler for facelifts.

[0175] The scaffold is usually delivered surgically, i.e., implanted into the body. It can be used for cartilage or osteochondral lesions. In the case of cartilage-related conditions, the scaffold is implanted at the debrided lesion site. The scaffold can be fixed to the lesion bed by absorbable sutures or fibrin glue. BRIEF DESCRIPTION OF THE DRAWINGS

[0176] Figure 1. In the creep test, the constant modulus (a), memory value (b), and permeability (c) were obtained under wet conditions (Example 3).

[0177] Figure 2. Dynamic (a) and static (b) slope moduli in the dynamic strain sweep test under wet conditions (Example 3).

[0178] Figure 3 . Dynamic modulus vs. test temperature under dry conditions. As can be seen, temperature has no effect on the modulus value, but there is a statistically significant (p = 0.00) difference between the non-sterilized and sterilized samples (Example 3). DETAILED DESCRIPTION

[0179] The present invention is further illustrated in the following non-limiting examples.

[0180] EXAMPLES

[0181] Example 1

[0182] A method for preparing a sterilized scaffold for medical use using poly(lactide) fibers and collagen, a method for drying the structure by freeze-drying to obtain the scaffold, and a method for sterilizing the scaffold to obtain a sterilized and stabilized PLA-collagen scaffold rhCo-PLA. In this example, it was demonstrated that the PLA component was not negatively affected by the processing method.

[0183] The scaffold was manufactured as follows. A medical-grade poly(L / D)lactide 96 / 4 (Corbion, Purac Biochem bv, Gorinchem, the Netherlands) with an inherent viscosity range of 1.8 dl / g to 2.2 dl / g and a residual monomer amount of less than 0.1% was used to manufacture thin fibers by melt spinning. Before melt spinning, the PLA raw material was dried in a vacuum oven. The melt processing of the fibers was carried out by melt spinning in a protective atmosphere within a temperature range of 70 - 240 °C. The spinning device consisted of a micro extruder and a high-speed spinning machine. The fibers were cut into short fibers and carded into a web. The PLA felt was made by needling the carded PLA web. The PLA felt was washed and dried in a laminar flow hood and then packaged before being placed in a clean room environment. The formation of recombinant human type III collagen (FibroGen Ltd., CA, USA) fibrils was accomplished by raising the pH of the collagen solution to 7 with an alkaline buffer solution. After the collagen gel was formed, the PLA felt was fully immersed in the collagen gel and placed in a sample mold. Then these structures were freeze-dried to obtain a completely dry structure, namely the rhCo-PLA scaffold. The manufactured scaffold was crosslinked at room temperature (RT) 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)). Then the manufactured rhCo-PLA scaffold was washed, freeze-dried, and packaged, and then sent for sterilization by gamma irradiation ≤25 kGy and under dry ice to avoid temperature rise during the sterilization process.

[0184] For the manufactured rhCo-PLA, the following studies were carried out. The monomer amount of L-lactide was measured by GC-MS technology (Rambol Analytics, Lahti, Finland) with a lower limit of 0.01 wt%, and the inherent viscosity (i.v.) was measured with a Lauda PVS viscometer (Lauda DR.R. Wobster GmbH, KG, Kornisseehofen, Germany). Samples were prepared by dissolving the polymer in 1 mg / ml chloroform. The viscosity was determined using an Ubbelohde 0c type capillary viscometer ( Mainz, Germany).

[0185] By this manufacturing method, the sterilized scaffold rhCo-PLA for medical use with poly(lactic acid) fibers and collagen has the following characteristics as shown in Table 1:

[0186] The residual monomer amount of the PLA component has not changed significantly as the monomer amount remains below 0.1 wt%, the same as that of the raw material monomer. The processing temperature of the used PLA and the sterilization have not widely changed the inherent viscosity of the extruded PLA: the acceptable reduction in this experiment is about 50%.

[0187] Table 1.

[0188]

[0189] Example 2

[0190] A scaffold for medical use with poly(lactic acid) fibers and collagen is obtained by drying the structure via freeze-drying to obtain the scaffold and sterilizing the scaffold to obtain the sterilized PLA-collagen scaffold. The rhCo-PLA scaffold as described in Example 1 is sterilized by gamma irradiation and the sterilization effect, especially on the collagen component, is evaluated. In addition, different gamma irradiation doses (under dry ice) are used on the rhCo-PLA scaffold to show the effect of the sterilization dose.

[0191] Data analysis is carried out according to the procedures described in detail in US Patent 10379106B2. The analysis includes extracting invariant data such as viscous stiffness and memory values under static and dynamic conditions from the data of the applied load (stress) and deformation (strain) respectively. In this specification, the following definitions are used:

[0192] The "invariant modulus" is the inherent elastic modulus value that does not depend on time or frequency and can be used to predict the material behavior (i.e., the true value).

[0193] The "dynamic invariant modulus" is the ratio of the dynamic stress amplitude to the dynamic true (logarithmic) strain amplitude, represented by real (non-complex) algebra (different from the commonly used definitions of real (storage) and imaginary (loss) moduli)

[0194] The "memory value" is a non-time-varying property of the sample, whose value ranges from zero to one and represents the viscous tendency of the material, even if the material itself is not a fluid. The memory value has no theoretical prediction and always has to be determined from experiments. In the present invention, experimental measurements of the memory value are carried out separately for static (creep) and dynamic conditions as they are found to be different.

[0195] The "fluid mobility" is a measure (coefficient) of the rate of fluid movement in the porous body, similar to the diffusion coefficient (using the same unit, in mm 2 / s). However, its properties are different from the latter because the movement of the fluid is not only through diffusion, but also due to convective parts and momentum transfer. Fluid mobility describes the likelihood that a fluid will flow inward under specific conditions.

[0196] "Apparent permeability" describes the ability of a porous body to allow a fluid to pass through its porous network, and it is measured in square distance units (m 2 ). It is a quantitative measure of the 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 is different from the permeability usually defined by Darcy's law because the latter requires an increase in the fluid pressure gradient across the material sample. However, as shown in U.S. Patent 10,379,106 B2, the methods also applied here allow the measurement of permeability without knowing the fluid pressure gradient, and to distinguish these methods, the term "apparent permeability" is used (expressed in millidarcies; 1 millidarcy = 10 -15 m 2 ).

[0197] A first test was performed on rhCo-PLA produced aseptically (rhCo-PLA-A), which was compared with rhCo-PLA sterilized with a standard irradiation dose, where the irradiation dose ≥ 25 kGy. In this process, the actual irradiation dose was measured as 29 kGy (rhCo-PLA-S). The aseptically produced rhCo-PLA (rhCo-PLA-A) has a gamma-irradiated PLA component (sterilized with a standard irradiation dose ≥ 25 kGy), but the addition of the collagen component is carried out under aseptic conditions after sterilization, i.e., the collagen component itself is not irradiated. Therefore, the PLA fiber component should have the same properties and contribution to the overall biomechanical properties, and the differences are mainly due to the role of collagen and its treatment.

[0198] Subsequently, the manufactured scaffolds were tested using a biomechanical test procedure with dynamic mechanical analysis (DMA) in a standard compression sample container (diameter 15 mm) within a dynamic mechanical analyzer DMA242E (Netzsch GmbH, Selb, Germany).

[0199] One part of the scaffold was subjected to a creep test under a constant force of 0.2 N stress, and another part was subjected to an oscillating force at a frequency of 1 Hz, generating a strain in the range of 5 - 50 μm (strain-scanning method). Briefly, after bringing the probe into contact with the sample and balancing the offset, the initial height of the sample immersed in the medium was measured again and further used as the initial height for true strain calculation.

[0200] In all cases, the scaffolds were fully immersed in water at room temperature for testing and allowed to equilibrate for 15 minutes prior to measurement. Thus, all samples were fully impregnated and no bubbles or dry areas were observed. The cross-sectional area of the tested scaffolds was 20 - 26 mm 2 . All tests were conducted for up to 300 minutes (until the dimensional change approached a constant value; displacement resolution ±0.0005 μm). The data were stored in ASCII text files and exported to data processing software (Microsoft Excel with custom code). The raw data were converted to stress and true strain, and the ratio of strain to stress vs. experimental time. Thereafter, a numerical algorithm of time convolution was applied, and the processed data were integrated non-locally and pairwise, line by line, using the mathematical method described in detail in U.S. Patent 10,379,106 B2.

[0201] These experimental data are shown in Table 2.

[0202] Table 2.

[0203]

[0204]

[0205]

[0206] Table 2 indicates how the sterilization carried out using the standard procedure of gamma irradiation with a dose of ≥25 kGy affected the biomechanical properties of the rhCo-PLA scaffolds as follows:

[0207] In the creep test, when the sterilization dose was 29 kGy, the rhCo-PLA-S scaffolds underwent a decrease in the unchanging modulus, which decreased significantly (-22%) after irradiation. The memory value also decreased (-43%), indicating that the rhCo-PLA scaffolds became more elastic after sterilization. The decrease in the effective fluid mobility and apparent permeability of the sterilized rhCo-PLA scaffolds (nearly twice) indicates a state with less mobile fluid within the scaffolds, and thus the irradiated scaffold structure became less permeable to fluids. These changes are undesirable, and such poor biomechanical properties of rhCo-PLA-S are unacceptable. Therefore, for the creep test, this sterilization method was found to severely affect the biomechanical properties of the rhCo-PLA scaffolds, thus failing to ensure their use in the intended applications.

[0208] In the strain sweep test, the dynamic unchanging modulus increased (+33%), indicating that the rhCo-PLA-S scaffolds became stiffer after sterilization, resulting in a poorer ability to adapt to dynamic strains in the surrounding tissues.

[0209] The second step was to evaluate the effect of more precise gamma irradiation doses on the rhCo-PLA scaffolds. Studies were conducted at a standard ≥25 kGy to compare different doses of gamma irradiation on the rhCo-PLA scaffolds. Thus, the rhCo-PLA scaffolds were sterilized with low-dose gamma irradiation 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.

[0210] The fabricated scaffolds were tested using the dynamic mechanical analysis as described above in the biomechanical testing procedure, using the same dynamic mechanical analyzer DMA242E (Netzsch GmbH, Selb, Germany). Similar to above, the scaffolds were subjected to a creep test under a constant force of 0.2 N. In all cases, the scaffolds were tested with complete immersion in water at room temperature and allowed to equilibrate for 15 minutes before measurement. The area of the tested scaffolds was 37 to 46 mm 2 .

[0211] Table 3 shows the biomechanical test results of the rhCo-PLA scaffolds under these different lower doses of gamma irradiation. The results showed that these different amounts of gamma irradiation ≤25 kGy did not cause significant changes in the biomechanics of the rhCo-PLA scaffolds (this was regarded as, for example, the storage modulus value, and the memory value of the G0 sample was within the limits measured for the sterilized samples G18 - G25).

[0212] Table 3.

[0213]

[0214] Example 3

[0215] A scaffold for medical use having poly(lactide) fibers and collagen, by drying the structure via lyophilization to obtain a scaffold, and sterilizing the scaffold to obtain a sterilized PLA-collagen scaffold.

[0216] The rhCo-PLA scaffolds as described in Example 1 were sterilized with gamma irradiation at RT (S-RT) or at a lower temperature (-70 °C) (S-LT), and the effect of sterilization was evaluated and compared with non-sterile scaffolds (NS). In this test, possible changes in the biomechanical properties of the scaffolds were generally demonstrated by different sterilization methods at the same expected dose (25 kGy), and it was found whether the reduced temperature during sterilization had an impact on the biomechanical properties. As described below, biomechanical tests were performed on both dry and wet samples.

[0217] Biomechanical analyses were performed in the compression mode up to 60 °C under dry conditions and at 25 °C under wet immersion conditions. Using a biomechanical test procedure for dynamic mechanical analysis in a standard compression sample container (15 mm diameter) within a dynamic mechanical analyzer DMA242E (Netzsch GmbH, Selb, Germany), the fabricated scaffolds were tested. A portion of the scaffolds was subjected to a creep test under a constant force of 0.2 N, and another portion was subjected to a strain sweep with an amplitude range of 5 to 25 μm at a frequency of 1 Hz. For the latter, the loading cycle was repeated 10 times, similar to Example 2.

[0218] The scaffolds were fully immersed in water at 25 °C and allowed to equilibrate for 15 minutes before measurement. Thus, all samples were fully impregnated, and no bubbles or dry areas were observed. A group of dry samples was additionally tested for the purpose of evaluating the thermal stability of the mechanical properties after heating to 60 °C at a rate of 2 K / min at 1 Hz and an amplitude of 25 μm under dry conditions (air).

[0219] The cross-sectional area of the tested scaffolds was 30 - 40 mm 2 . Data analysis was performed according to the procedure described in US Patent 10379106B2, with the aim of separately extracting invariant data such as viscous stiffness and memory values under static and dynamic conditions. The data are shown in Figures 1 - 3. It is worth mentioning that under repeated dynamic loading, all samples gradually shrank during each load sequence cycle. Therefore, the integral (slope) values of the dynamic stress / strain ratio ("standard dynamic modulus") and the associated static stress / strain ratio ("standard static modulus") at 1 Hz were extracted to represent the values covering all load cycles.

[0220] As seen from Figure 1, in the creep test, there were significant differences in certain characteristic values between some sample types, while others did not (error bars = standard deviation under wet conditions). For example, compared with S-LT, S-RT samples had a lower creep modulus and a higher creep permeability. Thus, there was an effect of temperature control (lowering the temperature) during the sterilization procedure.

[0221] Table 4

[0222] Wet conditions (mean)

[0223]

[0224]

[0225] Under dynamic wet conditions at 25 °C and 1 Hz (Figure 2), there were no significant differences in the dynamic and static stress / strain ratios (slope modulus) between the samples (NS, S-RT, and S-LT).

[0226] Surprisingly, under dynamic drying conditions with a temperature increase to 60 °C at 1 Hz, there were significant differences in the dynamic and static stress / strain ratios between the samples, p = 0.00 (for a two-sided arranged t-test with a 95% confidence interval, p < 0.05 is statistically significant), indicating that the S-RT and S-LT scaffolds became stiffer after the sterilization process ( Figure 3 ). These differences existed between non-sterilized (NS) and sterilized samples (S-RT and S-LT), but not between those sterilized at room temperature (S-RT) and at reduced temperature (S-LT). The change in temperature during the dynamic test had no effect on the properties up to 60 °C. Additionally, the test temperature did not seem to affect the sample stiffness (i.e., the stress / strain ratio of the material) up to 60 °C, so it can be concluded that the observed differences were mainly due to the sterilization process, and different temperatures; room temperature (S-RT) or lower (S-LT) had no effect on the properties of these samples.

[0227] These results indicate that sterilization (at RT or lower temperature (LT)) generally stabilized the scaffold properties, which was only clearly visible in these tests in the dry state.

[0228] From Figure 3 and Table 5, it can be seen that temperature had no effect on the modulus values, but there were statistically significant (p = 0.00) differences between non-sterilized and sterilized scaffolds.

[0229] Table 5

[0230] Drying conditions (mean)

[0231]

[0232] Generally, sterilization is considered to have a poor effect on the material properties of scaffolds, but in this case, sterilization led to unexpected results: gamma irradiation seems to have a positive effect on stabilizing the biomechanical properties of the scaffolds.

[0233] In summary, this Example 3 demonstrated the following effects of the sterilization procedure on the material vs. non-sterile material (therefore, the change is poor - the change in properties is not desired, general - neutral effect, no statistically significant difference, or good - the change in properties is desired):

[0234]

[0235] As mentioned above, a desired improvement 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.

[0236] As can be seen from the above table, for the following biomechanical parameters, an increase is desired: unchanging creep modulus (wet condition), dynamic modulus (dry condition), and dynamic modulus vs. temperature (dry condition), while a decrease or no change is desired for the following biomechanical parameters: creep permeability (wet condition) and dynamic modulus (wet condition).

[0237] As mentioned above, the desired improvement is when the increase (or decrease) is 1% or greater, 2% or greater, 3% or greater, 4% or greater, 5% or greater, 6% or greater, 7% or greater, 8% or greater, 9% or greater, or 10% or greater; or when the 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%.

[0238] Therefore, it can be assumed that the gamma irradiation sterilization procedure at RT or lower temperature (LT) is a positive step for these types of scaffolds to kinetically obtain a more stable structure both in the dry state and in the wet state. In addition, compared with sterilization at RT, the results obtained by sterilization at a lower temperature are superior, especially as shown under wet creep conditions.

[0239] It has also been confirmed that the rhCo-PLA scaffold as described in Example 1 has an unexpected reactivity to gamma sterilization conditions (dose and temperature control), such as improved biomechanical stability, allowing for more precise control of mechanical properties and the desired material optimization vs. clinical requirements. This effect is unexpected because it is well known that irradiation weakens or even destroys many organic materials and polymers.

Claims

1. A method for preparing a sterilized scaffold for medical use, the method comprising the following steps: i) loading collagen onto a fiber mesh containing poly(lactide) polymer or copolymer (PLA) fibers 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 sterilized scaffold, wherein the sterilization is carried out by gamma irradiation with up to 27 kGy.

2. The method according to claim 1, wherein, the obtained sterilized scaffold has improved biomechanical properties compared to the non-sterilized scaffold, and wherein the improved biomechanical properties are represented as an increase in one or more biomechanical parameters or biomechanical characteristics.

3. The method according to claim 2, wherein, the one or more biomechanical parameters or biomechanical characteristics are selected from the invariant creep modulus when tested under wet conditions and the dynamic modulus when tested under dry conditions.

4. The method according to claim 1, wherein, the obtained sterilized scaffold has improved biomechanical properties compared to the non-sterilized scaffold, and wherein the improved biomechanical properties are represented as a decrease or no change in one or more biomechanical parameters or biomechanical characteristics.

5. The method according to claim 4, wherein, the biomechanical parameters or biomechanical characteristics are selected from creep permeability and dynamic modulus, both of which are tested under wet conditions.

6. The method according to any one of claims 2 to 5, wherein, the increase or decrease is 1% or greater, 2% or greater, 3% or greater, 4% or greater, 5% or greater, 6% or greater, 7% or greater, 8% or greater, 9% or greater, or 10% or greater.

7. The method according to claim 4 or 5, wherein, 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%.

8. The method according to claim 2, wherein, the improved biomechanical property is stiffness.

9. The method according to claim 1, wherein, the temperature of the PLA-collagen scaffold before sterilization is substantially the same as or higher than the temperature of the PLA-collagen scaffold during sterilization.

10. The method according to claim 1, wherein, the sterilization is carried out at a temperature in the range of -200 °C to 40 °C.

11. The method according to claim 1, wherein, the gamma irradiation dose is in the range of 10 kGy to 27 kGy.

12. The method according to claim 11, wherein, the gamma irradiation dose is at most 25 kGy.

13. The method according to claim 1, wherein, the loading in step i) is carried out with a gel of collagen.

14. The method according to claim 13, wherein, the concentration of collagen in the gel is 0.1% to 2.0% w / w.

15. The method according to claim 1, wherein, The scaffold obtained in step i) contains 5 to 25% w / w collagen, the percentage being based on the total amount of PLA and collagen.

16. The method according to claim 1, wherein, the collagen is recombinant collagen, tissue-derived collagen or a combination thereof.

17. The method according to claim 1, wherein, the mesh containing PLA fibers and used in step i) of claim 1 is obtained by: i) providing PLA in solid form ii) subjecting the PLA to a process to obtain PLA fibers, and iii) subjecting the obtained fibers to a process to obtain a fibrous mesh.

18. The method according to claim 17, wherein, step ii) is carried out by spinning, such as electrospinning or melt spinning.

19. The method according to claim 17 or 18, wherein, the mesh in step iii) of claim 18 is subjected to a process including carding or needling to obtain a 3D network.

20. The method according to claim 1, wherein, PLA is poly(lactic acid).

21. The method according to claim 1, wherein, another crosslinking step is carried out before sterilization.

22. The method according to claim 21, wherein, the collagen in the PLA-collagen is crosslinked.

23. A scaffold obtainable by a method as defined in any one of claims 1 to 22.

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