RU2826997C2 - Biodegradable two-layer matrix for preventing formation of post-surgical adhesions, particularly in hernioplasty - Google Patents

Abstract

FIELD: chemistry.

SUBSTANCE: group of inventions relates to biodegradable matrix (14) for preventing the formation of postoperative adhesions in the body of a mammal, comprising upper layer (16) made of a first biocompatible polymer material containing polylactic acid as a main component; and lower layer (18) made of a second biocompatible polymer material containing poly (glycolic acid-lactic acid) as a main component; wherein the content of polylactic acid in the first polymer material is higher than in the second polymer material; both layers (16, 18) are formed in the form of porous frames containing a spongy structure with interconnected pores of different sizes, upper layer (16) is hydrophilic and has a wetting angle less than 75° and lower layer (18) is hydrophobic and has a wetting angle of more than 90°, and also relates to a method of producing a matrix by a) obtaining a first mixture of I salt particles and a dissolved first polymer material containing polylactic acid as a main component; b) applying a first mixture I to a surface to form a first layer; c) obtaining a second mixture II of salt particles and a dissolved second polymer material containing poly(glycolic acid-lactic acid) as a main component, wherein the content of polylactic acid in the first polymer material is higher than in the second polymer material; d) applying a layer of the second mixture II obtained in step c) on top of the first layer obtained in step b); e) drying the obtained structure to obtain a two-layer biodegradable matrix (14) with lower layer (18) of a second polymer material and upper layer (16) of a first polymer material; and f) plasma treatment of matrix (14) using plasma of oxidised gas at temperature below 50 °C.

EFFECT: group of inventions provides hernia closure without adhesion formation.

16 cl, 1 dwg

Description

The present invention relates to a biodegradable bilayer matrix for preventing the formation of postoperative adhesions, in particular after the restoration of soft tissues as a result of an abdominal surgical operation, for example, during hernioplasty in the body of a mammal, as stated in claim 1 of the claims. In addition, the present invention relates to a method for producing a biodegradable matrix, as defined in claim 10 of the claims.

Any tissue injury is usually followed by healing, which is usually accompanied by the formation of collagen scar tissue. Any physical, chemical or radiation injury to biological tissue consisting of various cells, extracellular matrix and connective tissue can lead to the death of these cells and structures. The healing process of damaged tissues includes various stages, which are also found in inflammatory processes. In particular, the first stages of wound healing include the removal of necrotic tissue, dead cells and cellular debris. In the second stage, the removed tissues are replaced, among other things, by inflammatory cells and fibroblasts, which are responsible for the production of collagen in its various forms, which leads to the formation of scar tissue. Such scar formation is necessary for the closure of a soft tissue defect. However, when two or more tissues in close proximity to each other are damaged, the scarring process can also lead to the formation of an undesirable connection of the originally separate tissues. Such a connection is usually called soldering.

In surgical procedures, multiple tissue trauma usually occurs prematurely and/or is caused by the incision made by the surgeon to reach the surgical site. As a result, postoperative adhesions between tissues and/or organs are one of the most common complications that occur after any type of surgery. Events such as excessive bleeding and/or inflammation or close contact between tissues significantly increase the likelihood of adhesions forming at the site of injury. The consequences for patients caused by such adhesions are often chronic pain and functional impairment, and in many cases, such consequences require reoperation.

Adhesions can have different shapes and strengths. In the field of abdominal surgery, adhesion formation is of particular concern. Adhesions in the abdominal region often result from abdominal trauma, which often occurs in motor vehicle accidents, or they occur as a result of surgical procedures. With regard to the latter cause, it has been found that intra-abdominal adhesions often occur at the site of surgical or interventional procedures in which the abdominal wall must be opened through all abdominal layers and the peritoneum. After the incision is sutured, the scarring process affects all abdominal layers, in particular also the peritoneal area along the incision. As a result, the physiological healing processes in the abdominal cavity in most cases cause the formation of an adhesion between the peritoneal tissues and the adjacent intra-abdominal organs. It has been shown that such adhesions occur in the form of dense adhesions forming wide constrictions or in the form of strong, taut veins several millimeters in diameter that can cross the entire abdominal cavity. It has even been found that extreme forms of abdominal adhesions cover entire areas of the abdominal cavity, making it impossible to separate, for example, loops of the small intestine and peritoneum and often leading to perforation of loops of the small or large intestine during exploratory surgery.

Although generally undesirable, the occurrence of postoperative adhesions is of particular concern in hernia repair surgery. A hernia is a condition defined as an abnormal displacement of an organ or tissue protruding through the wall of a cavity. Although hernias can occur in a variety of locations, they most commonly involve the abdominal cavity, specifically the abdominal wall, with the most common location being the groin. Inguinal hernias are most commonly referred to as inguinal hernias, but can also be femoral. Abdominal hernias occur in specific locations within vulnerable areas of the abdominal wall, such as a hiatal hernia, umbilical hernia, paraumbilical hernia, or Spiegel hernia, but they can actually occur anywhere in the abdominal wall. A special type is the incisional hernia, which is a hernia that occurs at the site of a previous surgical incision or interventional procedure.

Traditionally, hernias are repaired by open hernioplasty using sutures. Over the course of several months following hernia surgery, scar tissue gradually accumulates in the repaired area, so that the hernia defect closes and the abdominal wall is strengthened.

Unfortunately, it has been shown that in some patients the process of scar tissue formation is disrupted, which leads to the formation of another hernia after hernioplasty, i.e. to the so-called return or recurrence of the hernia. Therefore, particularly for the plastic surgery of larger hernias or in the case of a recurrence of the hernia, mesh implants are now widely used to reconstruct and strengthen the abdominal wall. Currently, commercially available meshes used in surgical reconstruction of soft tissues are either non-degradable or they are completely degradable and are absorbed by the patient's body after a certain time. Completely degradable meshes (often called "biological" meshes) were developed in the hope that they prevent the formation of adhesions.

Unfortunately, despite their usefulness in providing additional stability to the reconstructed area of the former soft tissue defect, both non-degradable and biodegradable meshes have been shown to cause post-operative adhesion formation. In particular, when using so-called “IPOM” meshes, i.e. “intra-abdominal pre-peritoneal overlay meshes”, to cover abdominal wall defects, it has been shown that the formation of unwanted inflammatory or fibrous bands or collagen scars connecting the mesh to the intra-abdominal structures is the cause of recurrent pain and obstruction of the small or large intestine. If the latter condition is not diagnosed in time, it can cause bowel necrosis, which usually requires more extensive surgeries and high-risk procedures. For this reason, in the case of abdominal hernia repair, the formation of post-operative adhesions can have a particularly serious impact on the patient’s health.

In order to prevent or at least minimize the formation of post-surgical adhesions, attempts have been made to isolate the injured tissue and separate it from any adjacent tissue using a biocompatible material. As a result, anti-adhesion barriers have been developed, which are now available in the form of physical films, fabrics, gels, or other materials that are applied between the layers of tissue at the end of a surgical procedure before the incision site is closed. Once in place, the anti-adhesion barrier acts as a physical barrier to separate the surfaces of the injured tissue, thereby preventing fibrin formation between the healing surfaces. Examples of commercially available anti-adhesion barriers for use in surgical procedures include, for example:

- Preclude® is a thin sheet of porous ePTFE (expanded polytetrafluoroethylene; also called GoreTex). It is a non-stick, microporous insert that is biocompatible and non-inflammatory. However, Preclude® is not resorbable or biodegradable, so additional surgery is required to remove it. It also must be sutured to the tissue in place. For this reason, Preclude® has not been approved in the United States for the prevention of adhesions.

- Seprafilm® (manufactured by Genzyme) is a clear, adhesive film composed of sodium hyaluronic acid and carboxymethylcellulose (CMC). Seprafilm® adheres to the tissues to which it is applied and is slowly absorbed by the body over seven days. Seprafilm® is approved for use in certain types of pelvic or abdominal surgery.

Interceed® (manufactured by Johnson & Johnson) is a knitted fabric composed of modified cellulose that swells and eventually gels when placed over the injured area and, like Seprafilm, forms a barrier and then slowly dissolves over several days. Interceed® is approved for use in pelvic surgery.

However, most of the currently available anti-adhesion barriers do not completely prevent adhesion formation. In particular, they are not suitable for use in hernia repair surgery, since in this case it is not enough to prevent the formation of connective tissue between the healing abdominal wall and the underlying intra-abdominal organs, but it is at least as important to ensure sufficient tissue growth from the adjacent tissues into the hernia area so that the gap in the abdominal wall is reliably closed. In particular, when using biodegradable (temporary) meshes or scaffolds for hernia closure, it is essential to induce fibrogenesis and scar formation, i.e. processes involved in wound healing, in order to form a stable scar plaque that prevents the formation of a recurrent hernia. Therefore, for successful hernia repair, agents are needed that induce fibrinogenesis and the formation of bridging tissue, which allows the soft tissue defect to be closed, but does not induce adhesion formation.

Thus, the problem solved by the present invention consists in obtaining a biodegradable matrix that ensures rapid, safe and stable closure of a soft tissue defect, in particular a hernia defect, while simultaneously reducing adhesion formation. At the same time, said matrix should be simple and inexpensive to manufacture and ensure the possibility of its use in traditional open and laparoscopic surgical methods.

The said problem is solved by means of a matrix according to claim 1 of the invention formula. Preferred embodiments are the subject of dependent claims of the invention formula.

According to the present invention, a biodegradable matrix is proposed for preventing the formation of postoperative adhesions during surgical restoration of soft tissues, in particular as a result of abdominal surgery in the body of a mammal. More specifically, the matrix according to the present invention is particularly well suited for covering areas of the peritoneum or abdominal wall that have been damaged during abdominal surgery or due to a hernia.

In the present application, the term "matrix" refers to a three-dimensional support, such as a mesh or scaffold, with a sponge-like structure suitable for cell colonization. In particular, the matrix according to the present invention has a sponge-like structure with interconnected pores of different sizes. In this sense, said matrix serves as a three-dimensional base that can be colonized with cells or tissue. Such colonization can occur in vitro or in vivo. Furthermore, in the context of transplants, said matrix serves to determine the location of the graft, as well as a placeholder marker for the tissue that is gradually formed in vivo.

The term "biodegradable" refers to a material that can be converted into metabolizable products in living organisms (or body fluids or cell cultures derived from living organisms). Biodegradable materials include, for example, polymers that are bioresorbable and/or biodegradable. "Biodegradable" means the ability to dissolve or suspend in biological fluids. Bioresorbable means the ability to be absorbed by the cells, tissues, or fluids of a living organism.

According to the present invention, the proposed matrix comprises a porous top layer made of a first biocompatible polymeric material comprising polylactic acid (generally abbreviated as "PLA") as a main component and optionally at least one additional polymer selected from the group consisting of polyglycolic acid (generally abbreviated as "PGA"), poly(glycolic acid-lactic acid), (generally abbreviated as "PLGA") and mixtures thereof. Thus, the term "main component" means that the content of PLA in the first polymeric material is higher than the content of any additional polymer that may be present in the first polymeric material.

The proposed matrix further comprises a porous lower layer made of a second biocompatible polymeric material containing as a main component at least one polymer selected from the group comprising polylactic acid [PLA], polyglycolic acid [PGA], poly(glycolic acid-lactic acid) [PLGA] and mixtures thereof, wherein the content of polylactic acid in the first polymeric material is higher than in the second polymeric material. Thus, the term "mixtures thereof" also includes copolymers of the named polymers.

As "biocompatible polymers", the polymers must be biologically tolerable and not cause rejection when introduced into a living organism. In the case of the present invention, biocompatible polymers also include polymers that the host recognizes as foreign, but whose rejection can be suppressed by the use of appropriate immunosuppression.

Both layers of the matrix are obtained in the form of porous frameworks having a sponge-like structure with interconnected pores of different sizes. In this regard, the term "porous" refers to a structure containing pores, i.e. cavities or empty areas. Such pores can have a round shape and/or an angular shape in a 2-dimensional section and/or an inclined shape when viewed in three dimensions. The shape of the pores can also be characterized by an expansion so that it can be compared with the shape of nerve cells. Although in general the term "pores" also refers to cavities formed by filaments surrounding empty areas, the pores within the meaning of the present invention are cavities formed in a sponge-like structure. Such cavities are thus limited by walls, as in natural sponges or corals. At least some of the pores or cavities are interconnected, which means that the pore walls between two adjacent pores can contain openings that form a connection between said adjacent pores. This is in contrast to a knitted mesh structure. Thus, at least some of the pores are interconnected so that they share space to form a fluidically connected intercellular network. In this way, the cells can spread along the matrix structure. In this regard, it should be noted that the pores formed within the upper layer and the pores in the lower layer may differ structurally from each other, for example with respect to their shape, size and/or interconnectivity.

One of the key elements of the present invention is that both layers are biodegradable and have the form of porous scaffolds, wherein the upper layer is hydrophilic and has a contact angle of less than 75°, preferably less than 60°, and the lower layer is hydrophobic and has a contact angle of more than 90°.

The term "wetting angle" as used in the context of the present application refers to the wetting angle of water on a surface, i.e. the angle formed at the interface where the water comes into contact with the surface. Thus, the "water" used to measure the wetting angle refers to pure water, in particular ultra-high purity water. In particular, the measurement of the wetting angle is carried out using the sessile drop method (e.g. using a device such as the EasyDrop DSA20E, GmbH), using a droplet size of 0.3 or 0.1 µl. In general, contact angles are calculated by fitting a circle segment function to the contour of a droplet placed on a surface (the "circle fitting" method).

The term "hydrophilic" or "hydrophilicity" as used in the context of the present invention refers to a contact angle of a surface area on a matrix that is less than 75°.

On the other hand, the term "hydrophobic" or hydrophobicity should be understood as a substrate having a surface area with a contact angle of more than 90°. As for the hydrophobic lower layer of the matrix according to the present invention, it preferably has a contact angle of more than 120°. Hydrophobic properties are usually a problem for synthetic polymers such as PLA; PGA and PLGA, and they are often enhanced by many post-treatment methods such as UV treatment.

It is noteworthy that the upper layer of the matrix according to the present invention is hydrophilic, despite the fact that it contains the above-mentioned mainly hydrophobic biodegradable synthetic polymers. Hydrophilic properties are important for facilitating cell penetration and adhesion.

It has been found that, with the first polymer material containing PLA as the main component, the hydrophilic properties of the top layer can be improved by applying certain post-treatment methods, in particular a low-temperature, low-pressure plasma treatment step, which will be described in more detail below. The term "plasma" here generally refers to an excited and radical-containing gas, i.e. an electrically conductive process gas containing electrons and ions. Plasma is usually generated by electrodes in a vacuum chamber (the so-called "high-frequency plasma method"), but it can also be produced by capacitive or inductive methods or microwave radiation. More detailed information on this subject is also given below in the experimental section.

The matrix according to the present invention provides many beneficial effects: on the one hand, it has been found that the hydrophilic properties of the upper layer of the matrix promote the growth of tissue, such as peritoneal cells, smooth muscle cells and fibroblasts, from the adjacent tissue into the wound to be closed, such as a hernia, and provide a uniform distribution of cells on the upper layer and throughout the upper layer. In addition, the porous nature of the upper layer provides the cells with a growth-stimulating environment, which promotes the formation of extracellular matrix tissue and various types of collagen fibers and, thus, the formation of a scar plaque that closes the tissue defect. Firstly, the growing scar plaque will create a strong connection between the degradable (and therefore temporary) matrix and the edges of the tissue defect. In the case of a hernia, the scar plaque is formed throughout the matrix, which allows the gap in the abdominal wall to be closed. Over time, as the matrix continues to degrade, the newly formed scar tissue will gradually take over the necessary support function by creating additional scarring, thereby preventing the wound from reopening or recurrent hernia formation.

On the other hand, it was found that the hydrophobic properties of the lower layer of the matrix effectively prevent the attachment of most types of cells, in particular inflammatory proteins or hydrophilic fluids, to the lower surface of the matrix, which prevents unwanted adhesion formation in tissues. In particular, when studying the plastic surgery of abdominal hernia, the hydrophobic properties of the lower layer of the matrix according to the present invention successfully prevented the infiltration of peritoneal or other physiological fluid into the matrix and minimized the unwanted formation of adhesions in tissues between the hernia plastic and intra-abdominal structures, in particular the small intestine.

Thus, the matrix according to the present invention provides temporary closure of a soft tissue defect, such as an abdominal hernia, and has the following advantages:

- On the one hand, the porous hydrophilic top layer facing the hernia promotes the ingrowth and proliferation of cells such as muscle cells and fibroblasts, which form new scar tissue that takes over the matrix's ever-decreasing support function. Eventually, after the matrix has completely degraded (i.e., when the polymer components of the matrix have been absorbed), there will be no permanent foreign material left in the patient's body.

- On the other hand, the hydrophobic properties of the lower layer of the matrix prevent the attachment of fibrin or inflammatory cell debris to the matrix from the intraperitoneal side (i.e. from the side of the matrix facing away from the hernia towards the abdominal cavity), so that the occurrence of inflammation and the growth of adhesive tissue between the matrix or newly formed scar plaque and the underlying tissues of the abdominal cavity can be avoided.

Although the matrix according to the present invention is particularly useful in hernia repair, it can be used to facilitate the healing process after surgery in general. For example, when performing abdominal surgery and there is intra-abdominal inflammation, it is important to provide a barrier between the inflamed intra-abdominal tissue and the overlying tissues that have been cut to gain access to the surgical area, regardless of the size of the wound, for example, even if only a small incision is made in a minimally invasive approach. Thus, the hydrophobic layer can provide this barrier, and the hydrophilic layer can promote the healing process of surgical wounds.

Another advantage of the matrix according to the present invention is that both layers are porous, which allows the use of pharmacologically active substances, such as epidermal growth factor, platelet growth factor, transforming growth factor beta, angiogenesis factor, antibiotics, antifungals, spermicides, hormones, enzymes and/or enzyme inhibitors, introduced into said layers, preferably into the upper layer, for delivering the listed substances to the wound site and positively influencing the growth of cells, such as collagen types IV and V, fibronectin, laminin, hyaluronic acid and proteoglycans, in the region of the upper layer and near it.

In order to facilitate the attachment and growth of cells throughout the top layer, it is preferable that the entire framework structure of the top layer has a hydrophilic surface, i.e. a surface whose contact angle is less than 75°. Such hydrophilic properties were achieved by plasma treatment of the matrix at a temperature below 50°C and preferably at a low pressure in the range of 10 ^-2 to 10 ^-6 bar, preferably in the range of 0.1 to 1.0 mbar.

As has been shown, since an increase in hydrophilicity correlates with improved cell attachment and proliferation, the contact angle of the matrix top layer is preferably less than 60°, more preferably less than 45°, and even more preferably less than 25°. Most preferably, the contact angle of the hydrophilic surface of the matrix top layer is from 0° to 10°, which means that such a top layer is "superhydrophilic".

The top and bottom layers can be designed as one integral structure, i.e. in which the two layers are manufactured as a single unit or are firmly connected to each other in several places. Alternatively, the two layers can also be designed as two separate structures that are separated from each other. One example would be a structure in which both the top layer and the bottom layer are sheet-shaped and simply laid loosely on top of each other.

According to a particularly preferred embodiment, the material compositions of the two layers differ from each other. The difference may consist in the type of polymer and/or in the content of a particular polymer in the polymer material. For example, both layers may consist of polymers of the same type, but with different polymer ratios. Alternatively, the upper layer and the lower layer may differ with respect to the type of polymers present in the polymer material.

According to one preferred embodiment, the first polymeric material of the top layer consists of at least 70% PL. (It is noteworthy that the term "PLA" includes all chiral forms of PLA, i.e. PLLA, PDLA and mixtures (copolymers) thereof.) This means that the top layer can be formed entirely of PLA or can consist of 70% or more PLA and 30% or less of at least one additional other polymer, such as PLG. When at least one additional polymer is present, the PLA and the other polymer(s) can form a copolymer, or the top layer can be formed as two separate components, such as a PLA base structure with a coating of other polymer(s). PLA has been found to have good tensile strength and a high elastic modulus. In addition, PLA has been found to be useful in providing and maintaining hydrophilic properties. More specifically, it has been found that for porous structures made of a material with a high PLA content, for example, if the first polymer consists of at least 70% PLA, the surface hydrophilicity can not only be significantly increased by plasma treatment as described herein, but said hydrophilicity can also be maintained for a long time. In fact, the hydrophilicity can also be maintained after sterilization of the matrix with hydrogen peroxide (as will be further described below). Due to the hydrophilic properties, the top layer facilitates the attachment of cells to the matrix and their ingrowth into it.

However, pure PLA has the disadvantage that it is less stable, i.e. less resistant to deformation, than, for example, PGA. However, since the bottom layer can be used to provide additional stability to the matrix, the top layer can also essentially consist of PLA. A specific example of a preferred PLA material is poly(L-lactide), which is available from Sigma Corporation (PLLA catalog number P1566), with a molecular weight of 85,000 to 160,000 Da. A suitable alternative is PLLA from Durect Corporation (Lactel® catalog number B6002-2).

As mentioned, the top layer polymer material can be based on PLA, but in combination with one or more other polymer(s) to enhance the stability of the top layer. One of the preferred additional polymers is PGA. Copolymers of PLA and PGA, the so-called "poly(lactide-co-glycolic acid)" (abbreviated PLGA or PLG), are available in various PLA/PGA ratios with well-defined physical properties. By varying the ratio of PLA to PGA copolymers, the different PLGA copolymers provide a wide range of elasticity and variable degradation rates ranging from a few days to several years.

According to a preferred embodiment, the second polymeric material of the lower layer of the matrix preferably consists of poly(glycolic acid-lactic acid) [PLGA]. In general, the higher the proportion of PGA in the PLGA composition, the higher the stability of the polymer. In addition, the higher the PGA content of the second polymeric material, the higher the hydrophobicity of the lower layer, even after plasma treatment is used to increase the hydrophilicity of the upper layer. Again, the lower layer generally provides additional stability to the matrix, while the upper layer provides a hydrophilic, cell-friendly environment that enhances the survival and proliferation rate of cells on and within the upper layer. As such, the second polymeric material of the lower layer will generally have a higher PGA content than the first polymeric material of the upper layer (since the first polymeric material of the upper layer has a higher PLA content than the second polymeric material of the lower layer).

According to one embodiment, both layers consist of PLGA, but with different ratios of PLA and PGA. A preferred polymer material for the top layer is an 85:15 mixture of poly(L-lactic acid) (PLLA) and PGA, i.e. a polymer mixture with a lactic acid (PLA) content of about 85 mol % and a glycolic acid (PGA) content of about 15 mol %. Such an 85:15 mixture can be purchased, for example, from Evonik Industries AG (Essen, Germany) or from Durect (Cupertino, California, USA) under the trade name RESOMER® RG 858 or LACTEL® Absorbable Polymers. The D,L-lactide-glycolide copolymer in the lower layer may be a 50:50 blend of PLLA and PLG, such as RESOMER® RG 502. Other preferred polymer blends for the upper and/or lower layer are a 65:35 D,L-lactide-glycolide copolymer, such as RESOMER® RG 653; a 75:25 D,L-lactide-glycolide copolymer, such as RESOMER® RG 752; a D,L-lactide-glycolide copolymer (always selected such that the PLA content in the first polymer material is higher than the PLA content in the second polymer material).

Methods for producing porous networks from the above-mentioned synthetic polymers are well known in the art. One possibility is to use the salt leaching technology described, for example, in EP 2256155.

It is preferable that the upper layer comprises or is at least partially coated with at least one natural polymer selected from the group consisting of collagen, gelatin, laminin, fibrinogen, albumin, chitin, chitosan, agarose, hyaluronic acid alginate and mixtures thereof, with collagen being preferred. The natural polymer imparts additional stability, hydrophilicity to the upper layer and promotes cell proliferation. The porous framework of the upper layer is preferably coated with the natural polymer or coated with the natural polymer in such a way that the underlying porous structure of the upper layer does not change upon coating. More specifically, it is preferable that the natural polymer coats the surface of the sponge-like structure without forming additional three-dimensional structures in the pores of the sponge-like structure of the upper layer. This ensures that said coating does not adversely affect the ability of cells to penetrate and spread in the upper layer of the matrix.

According to a particularly preferred embodiment, the second polymeric material of the lower layer essentially consists of PLGA or PGA, and the first polymeric material of the upper layer of the matrix essentially consists only of PLA (which can be coated with a natural polymer selected from the polymers listed in the previous paragraph). Of the mentioned natural polymers, collagen is the most preferred. This is due to the fact that collagen is a biomolecule of the extracellular matrix (ECM) and is the main component of the skin and bone. Due to its nanofibrous architecture, it is particularly effective in terms of stimulating cellular adhesion, growth and differentiated function in tissue cultures. However, it has also been found that the presence of collagen in the first polymeric material particularly enhances the hydrophilic properties of the upper layer of the matrix.

It is noteworthy that the term "collagen" used in the context of the present invention includes collagens of natural origin and synthetically produced collagens, as well as substances obtained from collagen, such as gelatin, which is a hydrolyzed form of collagen. In addition, the term "collagen" also includes all types of collagen. For example, a natural polymer can include only one specific type of collagen, such as type I, or can consist of a mixture of different types of collagen, such as a mixture of collagen type I and collagen type IV. In the latter case, preference is given to a mixture containing proteins in about percent by weight. Collagen type I is most preferred because it is one of the main components of natural blood vessels and provides a secondary structure with sites of cell attachment, as well as tensile strength. In addition, it is one of the main components of natural blood vessels and provides a natural site of attachment of cells involved in the wound healing process. Last but not least, the degradation product of type I-III collagen has also been shown to induce chemotactic fusion of human fibroblasts, which is particularly useful for the intended use of the matrix according to the present invention in soft tissue surgical repair.

According to a preferred embodiment, at least one layer of the upper layer and the lower layer, preferably both layers, has/have a flat sheet-like shape and is/are elastically deformable, allowing it/them to be folded or rolled up. In particular, it is preferred that the entire matrix is elastically deformable, so that it can be folded or rolled up and at the same time it can return to its original shape. This allows the matrix to be introduced, for example, through a trocar in a laparoscopic procedure, which ensures IPOM introduction of the matrix.

In general, the thickness of each layer is preferably at least from 0.1 mm to 20 mm, more preferably from about 1 mm to about 10 mm, even more preferably from about 1 mm to about 3 mm. It goes without saying that the two layers mentioned may also have different thicknesses.

When obtained in sheet form, the external shape (as viewed from above or in longitudinal section) of the matrix may be of any kind, such as rectangular, square, round, oval, etc., and said matrix may also be cut to match the shape of the soft tissue defect to be repaired. The external cross-sectional shape is preferably round or oval to avoid any sharp edges.

Furthermore, it is preferable that the porosity of the upper and lower layers - and preferably the entire matrix - is at least 80%, preferably at least 85%, more preferably at least 90%. Such porosity ensures that nutrients can diffuse through the matrix, creating a cell-friendly environment in the hydrophilic upper layer, which promotes cell proliferation and development. In addition, the porous structure allows the inclusion of growth factors or other molecules that stimulate cell growth in the matrix, in particular in the upper layer.

With regard to the degradation time of the matrix in the body, which usually occurs as a result of bioabsorption of its components, it is preferable that the upper layer has a higher degradation rate than the lower layer. In particular, the degradation time of the upper layer in the body is preferably inextricably linked to the formation of scar tissue, which reliably closes the soft tissue defect during matrix degradation. In particular, in hernia repair, it is highly preferable that the lower layer still provides additional support at the time of complete degradation of the upper layer, since this prevents the reopening of the former hernia until the scar plaque that forms on top of and within the degrading upper layer becomes strong enough to withstand the pressure in the abdominal cavity. In addition, the hydrophobicity of the lower layer helps to create a physical barrier between the forming scar tissue and the underlying abdominal organs. Therefore, during the first few months, when the growth of such cells is most significant, the formation of adhesions between the new scar tissue and the intra-abdominal organs is effectively prevented due to the hydrophobic properties of the lower layer. By the time the formation of a stable scar plaque covering the former hernia is complete, the underlying layer will continue to degrade, so that in general, 12 to 24 months after matrix implantation, there will be no foreign material left in the body.

According to a preferred embodiment, the total degradation time of the proposed matrix in a living organism is less than 24 months, with the preferred degradation time of the upper layer being less than 6 months, preferably less than 4 months, and the preferred degradation time of the lower layer being at least 4 months, preferably from 4 to 24 months. Thus, the lower layer will provide an additional support function during the first 4 to 12 months. In general, after 24 months, the lower layer will also be more or less completely destroyed.

According to a specific embodiment, it is preferable that the degradation time of the upper layer in a living organism is from 1 to 4 months, preferably about 3 months. On the other hand, in the case of the lower layer, it is preferable that the degradation time in a living organism is from 6 to 12 months.

Another advantage of the matrix according to the present invention is that it allows the inclusion in the matrix, in particular in the upper layer, of substances that are subsequently delivered to the soft tissue defect. Preferred substances are collagen types IV and V, fibronectin, laminin, hyaluronic acid and proteoglycans. Likewise, pharmacologically active substances such as growth factors, antibiotics, antifungals, spermicides, hormones, enzymes and/or enzyme inhibitors can be introduced into the matrix.

In order to facilitate the formation of scar tissue in and around the top layer, it is preferable that the top layer additionally contains growth factors. Growth factors typically act as signaling molecules between cells and often promote cell differentiation and maturation. For example, epidermal growth factor (EGF) enhances osteogenic differentiation, while fibroblast growth factors (FGF) and vascular endothelial growth factors (VEGF) stimulate blood vessel differentiation (angiogenesis). In view of the use of the matrix for soft tissue repair, the top layer preferably contains at least one growth factor selected from the group consisting of interleukins, acidic fibroblast growth factor, basic fibroblast growth factor (b-FGF), epidermal growth factor, insulin-like growth factor, insulin-like growth factor binding protein, platelet-derived growth factor (PDGF), transforming growth factor alpha, transforming growth factor beta, VEGF and hepatocyte growth factor (HGF). Such growth factors are important for regulating cell proliferation and differentiation, protein synthesis, and ECM (extracellular matrix) remodeling. In particular, b-FGF, PDGF, VEGF, and HGF have been shown to increase granulation, epithelialization, and capillary formation through the secretion of angiogenic cytokines. They have also been shown to inhibit neutrophil and macrophage migration to the wound site by secreting factors that suppress migration and both IL-1α and IL-1β, and secrete anti-inflammatory factors that prevent apoptosis and improve wound healing.

According to a particularly preferred embodiment, the top layer comprises growth factors added to the matrix, in particular to the top layer, in the form of a secretome derived from mesenchymal placental cells. It has been found that a commercially available secretome derived from (or at least containing) Wharton Jelly Stem Cell (CM-hWAO) cultured under hypoxic conditions is particularly effective in stimulating cell attachment to and ingrowth into the top layer of the matrix. Such a stem cell secretome can be purchased, for example, from the Institute of Stem Cells and Cancer (PT Kalbe Farma Tbk.).

To obtain a matrix, you can use a method that includes the following stages:

a) obtaining a first mixture I consisting of salt particles and a dissolved first polymeric material containing polylactic acid [PLA] as a main component and optionally at least one additional polymer selected from the group consisting of polyglycolic acid [PGA], poly(glycolic acid-lactic acid) [PLGA] and mixtures thereof;

b) applying the first mixture I to the surface to form the first layer;

c) obtaining a second mixture II consisting of salt particles and a dissolved second polymeric material containing as a main component at least one polymer selected from the group consisting of PGA, PLA, PLGA and mixtures thereof, wherein the PLA content in the first polymeric material is higher than in the second polymeric material;

d) applying a layer of the second mixture II, obtained in step c), on top of the first layer;

e) drying the resulting structure to obtain a two-layer biodegradable matrix comprising an upper layer of a first polymeric material and a lower layer of a second polymeric material; and

e) plasma treatment of the matrix using oxidized gas plasma at a temperature below 50°C.

It was found that plasma treatment enhanced the hydrophilic properties of the surface of the top layer without adversely affecting the stability or structural integrity of the matrix. On the other hand, plasma treatment of PGA structures was not found to enhance hydrophilicity. Due to the higher PLA content in the top layer compared to the bottom layer, plasma treatment was particularly effective in enhancing, as well as maintaining, the hydrophilic properties of the top layer, but it did not enhance or only slightly enhanced the hydrophilicity of the bottom layer.

The ionized gas plasma used for plasma processing is preferably selected from the group consisting of helium, argon, nitrogen, neon, silane, hydrogen, oxygen and mixtures thereof. Preferred gases for processing are hydrogen, oxygen and nitrogen, in particular oxygen.

More particularly, the plasma treatment preferably comprises low-temperature, low-pressure plasma treatment, in which the carrier grid is exposed to ionized gas plasma at i) a temperature below 50°, preferably below 40°C, ii) for at least 2 minutes, more preferably from 5 to 20 minutes, and iii) at a pressure in the range from 10 ^-2 to 10 ^-6 bar, preferably at a pressure in the range from 0.1 to 1.0 mbar.

Thus, the term "plasma" generally refers to an excited gas containing radicals, i.e. an electrically conductive process gas containing electrons and ions. Plasma is usually generated using electrodes in a vacuum chamber (the so-called "high-frequency plasma method"), but it can also be produced by capacitive or inductive methods or microwave radiation.

Instead of the listed steps a) - e), the matrix according to the present invention can also be manufactured using 3D printing, electrospinning and other methods known in the art for manufacturing polymer scaffolds.

The proposed method may further include a stage in which the porous framework of the upper layer formed from the first polymer material is coated with a natural polymer selected from the group consisting of collagen, gelatin, laminin, fibrinogen, albumin, chitin, chitosan, agarose, hyaluronic acid alginate and mixtures thereof, preferably collagen. Such a stage, if present, is preferably carried out before stage f) of plasma treatment.

Taking into account its subsequent use as an implant, the matrix proposed in the present invention is then (i.e. after step f)) preferably sterilized. For this purpose, it is preferable to use a special sterilization method that has been developed for this purpose. Such a sterilization method allows sterilization of tissues sensitive to heat and/or UV radiation, in particular polymer scaffolds, and is therefore not limited to the specific matrix described above, but is applicable to all types of (heat-sensitive) articles that need to be sterilized.

Nowadays, there is no doubt that sterilization is necessary for almost any device and article used in the medical field, such as instruments, all types of implants and any surgical aids. Theoretically, there are many sterilization methods, but not all of them are applicable to all substrates. Metallic substrates, such as, for example, a metal instrument or implant, can be sterilized by heat using steam. This method is usually carried out in a steam sterilizer (also called an autoclave), using steam, the temperature of which is usually higher than 120 ° C, under pressure. However, heat sterilization is not suitable if the sterilized article is sensitive to heat. In addition, the use of steam is not suitable for components that are subject to biodegradation and, therefore, to a certain extent, soluble in water. Therefore, biodegradable polymer substrates containing a heat-sensitive natural polymer, such as, for example, collagen, cannot be sterilized with hot steam without damaging the molecular structure of such a substrate.

Alternatively, the substrate can be sterilized with ethylene oxide gas or plasma sterilization. However, with regard to the use of ethylene oxide, this method has another drawback, namely, the need for relatively strict safety measures due to the high toxicity of the sterilizing agent.

Additional sterilization methods include radiation sterilization, in particular gamma sterilization or X-ray sterilization. On the other hand, the main disadvantage of such methods is that the hydrophilic surface properties of the substrate (in this case the top layer) are usually lost or at least significantly deteriorated due to the sterilization treatment.

Thus, for sterilization of the two-layer matrix according to the present invention described above, a method is required that avoids the use of heat, i.e., temperatures above 50°C, to maintain the three-dimensional polymer structure of the matrix. In addition, said method must ensure the maintenance of high hydrophilicity of the upper layer during and after the sterilization procedure.

In view of the above, it was also an additional object of the present invention to provide a simple method for thoroughly sterilizing the matrix without deteriorating the hydrophilicity of the top layer.

The following procedure has been found to satisfy all of the above requirements and is therefore particularly well suited for sterilizing sensitive substrates such as the bilayer matrix of the present invention. The proposed procedure comprises the steps

I. obtaining a two-layer matrix as described in this application, and

II. exposing the matrix to a medium containing hydrogen peroxide at a temperature below 50°C, preferably below 40°C; at a reduced pressure of 10 ^-6 to 10 ^-2 bar; and for at least 2 minutes.

The unexpected discovery that such low-temperature hydrogen peroxide sterilization can produce a sterile and hydrophilic biodegradable product opens up the possibility of using a simple sterilization process for sensitive materials without adversely affecting their structural integrity and hydrophilic properties. As an additional advantage, the new sterilization method is very simple and effective, since it does not require labor-intensive preparatory steps.

The hydrogen peroxide containing medium can be obtained either by treatment with _{H2O2 plasma} or by placing the substrate to be sterilized in a vacuum chamber together with a source of (usually liquid) hydrogen peroxide. The plasma treatment preferably comprises low pressure plasma treatment, in which the matrix is exposed to ionized gas plasma at a pressure in the range of ^10-2 to ^10-6 bar, preferably in the range of 0.1 to 20.0 mbar, and at a temperature below 50°C, preferably below 40°C, for at least 2 minutes. Alternatively, the matrix can be placed inside a vacuum chamber and by applying a pressure low enough to evaporate the hydrogen peroxide, the hydrogen peroxide evaporates, thereby creating an atmosphere containing hydrogen peroxide. Particularly preferred (negative) pressures are from 0.1 to 20.0 mbar, for example from 6 to 12 mbar.

_The sterilization time is strongly dependent on the pressure inside the chamber, the processing temperature and the concentration of _{the H2O2} solution. _{The H2O2} solution preferably contains _H2O2 in an amount of about 30% by volume or less. The preferred processing time is at least a few minutes, for example from 2 to 30 minutes, alternatively at least one hour. When using "higher" pressures, in particular above or about 10 mbar, and/or low temperatures, for example below 40°C, or _{an H2O2} concentration below 30%, for example from 20 to 25%, a processing time of several hours, for example from 10 to 12 hours, is preferred.

As mentioned earlier, the bilayer matrix according to the present invention is particularly useful for preventing the formation of intra-abdominal adhesions, for example in the area of hernioplasty, and for preventing the formation of a recurrent hernia. Thus, the present invention also relates to the use of the bilayer matrix proposed in the present invention in surgical restoration of soft tissues, in particular in a surgical operation on abdominal organs, for example in hernioplasty. In particular, when used for hernioplasty, the method according to the present invention comprises the steps of:

i. obtaining a biodegradable matrix having a hydrophilic top layer and a hydrophobic bottom layer as described in the above sections;

ii. cutting the skin and tissues of the patient's abdominal cavity to gain access to the hernia in the abdominal wall;

iii. placement of the matrix either over the defect, such as over the abdominal wall, or alternatively under the muscular layer of the abdominal wall (sublay technique) or as an intraperitoneal overlay mesh (IPOM) below the peritoneum;

iv. closure of the incision.

If necessary, the matrix can be additionally attached to the abdominal muscles to prevent movement.

Preferred embodiments with respect to the structure of the matrix and its placement in the patient's body during hernioplasty are further illustrated with the aid of the accompanying figure, wherein

Fig. 1 shows a schematic representation of an incision passing through a soft tissue defect, in particular a hernia, eliminated using a matrix in accordance with the present invention.

The schematic representation in Fig. 1 shows a defect (gap) 10 in muscle tissue 12, which has been bridged by means of a matrix 14 according to the present invention. In particular, a biodegradable matrix according to the present invention is shown, having a porous hydrophilic top layer 16 and a porous hydrophobic bottom layer 18. The hydrophilic top layer 16 consists of polylactic acid [PLA] and optionally collagen. Said top layer has a contact angle of less than 10° and is thus superhydrophilic. The bottom layer 18 of the matrix consists of poly(glycolic acid-lactic acid) [PLGA], for example Resomer® RG 503 H from Sigma Aldrich with a lactide:glycolide content of 50:50, and has a contact angle of more than 90°. Said bottom layer is thus hydrophobic. The two layers are bonded to each other along a common interface 20 and form a single two-layer matrix device 14. The hydrophilic properties of the upper layer are provided by plasma treatment of the matrix using gaseous oxygen plasma at a temperature below 40°C and a pressure in the range from 0.1 to 1.0 mbar. It has been shown that such plasma treatment allows obtaining an upper layer with a high PLA content, having hydrophilic properties, while the hydrophobic properties of the lower PLGA layer do not change significantly during plasma treatment.

In surgical hernia repair, an incision is made through the patient's skin 22 and the abdominal tissue 12 to gain access to the hernia 10 in the abdominal wall. The matrix 14 is then placed either above the defect, such as the abdominal wall (Lichtenstein inguinal hernia repair) or alternatively under the muscular layer of the abdominal wall (sublay method) or, as in the case shown in Fig. 1, in the form of an IPOM (intraperitoneal overlay mesh) below the peritoneum 24. The lower layer 18 will face the intra-abdominal organs, in particular the small intestine 26, and the upper layer will face the hernia 10. If desired, the matrix 14 can be additionally attached to the abdominal muscles 12 and/or the peritoneum 24 to prevent displacement. (Before the matrix is placed, the defect in the abdominal wall can also be closed with sutures. The matrix is then placed above or below the repaired defect to stabilize the suture material and facilitate the healing process.) The incision is then closed with sutures 28 by closing the overlying skin 22.

In general, the following processes occur after implantation:

As a degradable structure, the proposed matrix creates a temporary support structure for cells that can migrate into the matrix, in particular into the upper layer of the matrix, from adjacent tissues and proliferate. The hydrophilic properties of the upper layer facing the hernia promote the ingrowth and proliferation of cells, such as muscle cells and fibroblasts, which form new scar tissue that takes over the ever-decreasing support function of the matrix. On the other hand, the hydrophobic properties of the lower layer have the opposite effect: said layer prevents the attachment of cells, such as, among other things, inflammatory cells, fibrin or cellular debris, to the matrix from the intraperitoneal side, i.e. from the side of the matrix facing away from the hernia towards the abdominal cavity, so that the occurrence of inflammation and the growth of adhesion tissue between the matrix or the newly formed scar plaque and the underlying tissues of the abdominal cavity can be avoided.

The upper layer has a higher rate of degradation than the lower layer. The degradation of the upper layer takes approximately 3 to 6 months and is inextricably linked to the formation of scar tissue that reliably connects the edges of the former hernia to the abdominal wall. By the time the upper layer has completely degraded, the lower layer still continues to provide additional support against abdominal pressure and also provides a physical barrier between the newly formed scar tissue and the underlying abdominal organs. Therefore, at a time when cell growth is most significant, adhesions between the scar tissue and the intra-abdominal organs are effectively prevented by the hydrophobic properties of the lower layer. In addition, in the presence of intra-abdominal inflammation, this physical barrier also separates the inflamed tissue from the surgical wound. After approximately 12 months, when a stable scar plaque covering the former hernia has been completed, the lower layer will also usually be essentially completely degraded, leaving no foreign material in the body.

The matrix can be simply placed over the defect or additionally secured in place with sutures or other measures to prevent matrix displacement.

For several days and weeks after the matrix is injected, cells from the surrounding tissues, particularly smooth muscle cells and fibroblasts, will continue to proliferate and form scar tissue that firmly closes the hernia. By the time both layers of the matrix have completely disintegrated, this scar tissue will have reliably closed the hernia.

In order to prove the above-described effects, the matrix proposed in the present invention was examined in vivo by implantation in rats. In particular, an experimental abdominal wall hernia in a rat was repaired using a two-layer matrix according to the present invention. In such an experiment, the matrix was placed below the peritoneum, as shown in Fig. 1. In order to fix the implanted matrix in place, before closing the skin with sutures, said matrix was attached to the muscle tissue with several sutures. After six weeks of healing, the implantation site was reopened. It was found that no adhesions were formed between the matrix and the underlying small intestine. On the other hand, in the area that was not covered by the matrix, an adhesive constriction was formed from the edge of the liver to the abdominal wall. Thus, the in vivo experiment confirmed that the matrix according to the present invention successfully prevents the formation of postoperative adhesions.

In the following sections, specific examples of methods for producing a matrix according to the present invention will be described in detail.

Experimental data

Obtaining a matrix

Solid sodium chloride (NaCl) particles were ground using a mortar and pestle before sieving to obtain NaCl particles with sizes ranging from 355 to 425 μm. 9 g of NaCl particles were placed in a centrifuge tube and dried in a desiccator. The NaCl particles were then placed in an aluminum tray and a PLLA solution prepared from 1 g of PLLA beads (lactide 100 from Durect Lactel®) in 5 ml of chloroform was poured onto the NaCl particles. The PLLA solution was mixed with the NaCl particles and the mixture was then spread evenly in the aluminum tray to obtain a flat PLLA layer.

Additional post-treatment with collagen

Some scaffolds were also provided with a collagen coating deposited on the PLLA layer (which subsequently formed the top layer). For this purpose, the resulting PLLA layer was dried and separated from the aluminum tray. Then, a collagen solution (collagen type I solution; Wako) was poured into a Petri dish. The concentration of the collagen solution was chosen in the range of 0.1 to 5.0 (w/v)%, preferably 1% (w/v). Before placing in another Petri dish, the PLLA layer was immersed in the collagen solution and frozen in a low-temperature freezer at -80 °C for several hours before freeze-drying (also known as lyophilization) under a vacuum of <5 mbar (at a temperature of -50 °C to -80 °C) for at least 24 hours.

To form the hydrophobic PLGA layer (which would later form the bottom layer), 9.0 g of NaCl particles were placed in a second aluminum tray and a PLGA solution was prepared from 1 g of PLGA beads (lactide 25; glycolide 75 from Durect Lactel®) in 5 ml of chloroform. The PLGA solution was mixed with the NaCl particles. The PLGA/NaCl mixture was then poured on top of the PLLA layer (with or without collagen coating) prepared in the first aluminum tray and spread evenly to form a matrix with a PLGA layer on top of the PLLA layer.

The PLLA/PLGA-NaCl matrix was separated from the aluminum tray and dried in a vacuum chamber at -0.1 MPa for 3 to 4 days.

The resulting dried PLLA/PLGA-NaCl scaffolds were placed in a beaker, immersed in _ddH2O (double deionized water), and maintained in a linear shaking bath at 25°C (room temperature) at 60 rpm for 48 h to leach/wash out NaCl particles. The water in the beaker was changed every 1-2 h. The PLLA/PLGA bilayer scaffolds were removed from the beaker and dried in a fume hood overnight.

The resulting matrices contained pores with diameters ranging from 355 to 425 µm.

It was found that the PLGA layer could also be obtained first and the PLLA layer in the second step.

_O2 plasma treatment

The matrices, either with or without a collagen-coated top layer, were further subjected to plasma treatment using ionized gas plasma, preferably selected from the group consisting of helium, argon, nitrogen, neon, silane, hydrogen, oxygen and mixtures thereof. Preferably, plasma treatment using ionized oxygen gas plasma is used.

Plasma treatment was performed using a plasma treatment device from Diener (Diener electronics; Plasma-Surface-Technology; Ebhausen, Germany) in a vacuum chamber for a period of 5 to 20 minutes, preferably 8 to 15 minutes. The following treatment parameters were set: vacuum chamber pressure: 0.40 mbar; power: 35 W; gaseous oxygen flow rate: 5 ^cm3 /min (minimum) - 60 ^cm3 /min (maximum).

It was found that plasma treatment significantly increased the hydrophilic properties of the PLLA top layer, regardless of the presence of the collagen coating, but not of the PLGA layer. The latter remained hydrophobic.

Sterilization

The matrices, with or without a collagen coating on the top layer, were then sterilized by placing them in an _{H2O2 -} containing environment at a temperature below 40°C. The _{H2O2 -} containing environment was created inside a vacuum chamber by placing the matrix in the chamber together with an open flask or cup containing _{a H2O2} solution and then evacuating the chamber to evaporate _{the H2O2} . _{The H2O2} solution contained _H2O2 in an amount of 30% by volume or less. The processing time was strongly dependent on the pressure in the chamber and the concentration of _{the H2O2} solution. _The pressure inside the chamber was such that hydrogen peroxide evaporated. Preferred (negative) pressures were from ^10-2 to ^10-6 bar, preferably from 0.1 to 20.0 mbar, for example 9 mbar. The preferred treatment time was at least two minutes, such as 2 to 30 minutes. However, in particular at a pressure above 1 mbar, such as about 10 mbar, and/or at a lower temperature (such as below 35°C) and/or at a low hydrogen peroxide concentration (such as below 30 vol. %), the treatment time was preferably greater than one hour, such as 8 to 10 hours.

Measurements of static contact angle, sessile drop method

Contact angle measurements were performed to determine the degree of hydrophilicity or hydrophobicity. Typically, the contact angles of the top and bottom layers of the matrix were determined by measuring the static contact angle using the sessile drop method using ultrapure water (EasyDrop DSA20E, GmbH). The droplet size for contact angle measurements was set to 0.1 μl. Contact angles were calculated by fitting a circle segment function to the contour of a drop placed on the surface (circle fitting method).

Claims (29)

1. A biodegradable matrix (14) for preventing the formation of postoperative adhesions in the body of a mammal, containing a top layer (16) made of a first biocompatible polymeric material containing polylactic acid as a main component; and a lower layer (18) made of a second biocompatible polymeric material containing poly(glycolic acid-lactic acid) as its main component; wherein the content of polylactic acid in the first polymer material is higher than in the second polymer material; both layers (16, 18) are formed in the form of porous frameworks containing a sponge-like structure with interconnected pores of different sizes, the top layer (16) is hydrophilic and has a contact angle of less than 75° and The bottom layer (18) is hydrophobic and has a contact angle of more than 90°. 2. The matrix according to claim 1, wherein the first biocompatible polymeric material comprises at least one additional polymer selected from the group consisting of polyglycolic acid, poly(glycolic acid-lactic acid) and mixtures thereof. 3. The matrix according to one of claims 1, 2, in which the upper layer (16) has a wetting contact angle of less than 60°, preferably less than 45°, more preferably less than 25°, even more preferably less than 15°, most preferably in the range from 0° to 10°. 4. The matrix according to one of paragraphs. 1-3, in which the two specified layers (16, 18) are either formed as a single whole or are firmly connected to each other along a common interface (20). 5. A matrix according to one of paragraphs 1-3, in which the two layers (16, 18) are made in the form of two separate structures that are separated from each other. 6. The matrix according to one of claims 1 to 5, wherein the first polymeric material consists of at least 70%, preferably at least 80%, polylactic acid or consists essentially of polylactic acid. 7. The matrix according to one of claims 1-6, in which at least one, preferably each of the two layers (16, 18) has a flat sheet-like shape and is elastically deformable, which allows it to be folded or rolled up. 8. The matrix according to one of claims 1-7, wherein the top layer (16) comprises or is coated with at least one natural polymer selected from the group consisting of collagen, gelatin, laminin, fibrinogen, albumin, chitin, chitosan, agarose, hyaluronic acid alginate and mixtures thereof, wherein collagen is preferred. 9. The matrix according to one of paragraphs. 1-8, wherein the postoperative adhesions are formed after restoration of soft tissues as a result of abdominal surgery. 10. The method for obtaining the matrix according to paragraphs 1-9 by a) obtaining a first mixture of I salt particles and a dissolved first polymer material containing polylactic acid as a main component; b) applying the first mixture I to the surface to form the first layer; c) obtaining a second mixture of II salt particles and a dissolved second polymeric material containing poly(glycolic acid-lactic acid) as a main component, wherein the content of polylactic acid in the first polymer material is higher than in the second polymer material; d) applying a layer of the second mixture II obtained in step c) on top of the first layer obtained in step b); e) drying the resulting structure to obtain a two-layer biodegradable matrix (14) with a lower layer (18) made of a second polymeric material and an upper layer (16) made of a first polymeric material; and f) plasma treatment of the matrix (14) using oxidized gas plasma at a temperature below 50 °C. 11. The method of claim 10, wherein the first biocompatible polymeric material comprises at least one additional polymer selected from the group consisting of polyglycolic acid, poly(glycolic acid-lactic acid), and mixtures thereof. 12. The method according to one of claims 10, 11, wherein the first polymeric material consists of at least 70% polylactic acid. 13. The method according to one of claims 10-12, wherein the plasma treatment is carried out for 2 to 30 minutes, preferably 5 to 20 minutes, and at a pressure in the range of 10 ^-2 to 10 ^-6 bar, preferably in the range of 0.1 to 1.0 mbar. 14. The method according to one of claims 10-13, further comprising the step of coating the top layer (16) of the matrix (14) obtained in step e) with a natural polymer selected from the group consisting of collagen, gelatin, laminin, fibrinogen, albumin, chitin, chitosan, agarose, hyaluronic acid alginate and mixtures thereof, preferably collagen. 15. The method according to one of claims 10-14, further comprising the step g) of sterilizing the obtained matrix (14) by treating it with hydrogen peroxide at a temperature below 50 °C. 16. The method according to claim 15, wherein the treatment with hydrogen peroxide in step g) is carried out by exposing the matrix (14) to _{H2O2 plasma} or an atmosphere _{containing H2O2} under reduced pressure, preferably in the range of from ^10-2 to ^10-6 bar, more preferably in the range of from 0.1 to 20.0 mbar, and for at least 2 min, preferably from 2 to 30 min or alternatively for several hours.