WO2013091050A1

WO2013091050A1 - Method for producing polymer biomaterials using an infra-red laser

Info

Publication number: WO2013091050A1
Application number: PCT/BR2012/000502
Authority: WO
Inventors: Rubens MACIEL FILHO; Marcele Fonseca PASSOS; André Luiz Jardin MUNHOZ; Vanessa Petrilli BAVARESCO
Original assignee: Universidade Estadual De Campinas - Unicamp
Priority date: 2011-12-22
Filing date: 2012-12-06
Publication date: 2013-06-27
Also published as: BRPI1105297A2; BRPI1105297B1

Abstract

The present invention describes a method for producing biomaterials by using an infra-red laser to cause two polymer materials to adhere to each other. These biomaterials have numerous uses, in particular as polymer supports, scaffolds and coatings for artificial joint prostheses. The method improves the mechanical properties of the polymers, achieving satisfactory adhesion between the hydrogel and the substrate, minimising the wear of the components that form orthopedic devices, for example, which is one of the main causes of failure of such devices.

Description

PROCEDURE FOR OBTAINING VIA LASER POLYMERIC BIOMATERIALS

Field of the invention

The present invention relates to a process for obtaining polymeric biomaterials. More specifically, the present invention deals with a process of obtaining biomaterials by the adhesion of two polymers employing infrared laser.

The biomaterials obtained by the present invention may be used for cell support and growth, scaffolds, polymeric coatings for use in artificial joint prostheses.

Fundamentals of the invention

Synovial or diatrodial joints constitute the majority of the joints of the human body, making it possible to move around and perform daily activities and, therefore, when damaged, their restoration is of great importance. Depending on the level of joint involvement, its replacement by artificial components becomes indispensable and, in this case, arthroplasty, which may be total or partial, is performed. Prosthesis arthroplasty is indicated for patients who fail to treat degenerative diseases, such as osteonecrosis or osteoarthritis, or where they have progressed after joint preservation surgeries. Osteonecrosis presents a progressive clinical condition, evolving with limitation and functional incapacity due to pain, decreased range of motion, stiffness, and consequently muscle weakness.

The durability of these prostheses, however, does not exceed on average

15 years. This is because the useful life of its components depends on a number of factors, resulting in bone resorption. The presence of surface wear particles, present in the middle, generated by the lack of lubrication between the new articular surfaces, causes a response of the organism to this process, with consequent collapse between the bone-implant interface. Studies show that due to the wear process superficial, the percentage of prosthesis failure increases proportionally with the increased time of implant use (Bavaresco, 2000). In view of this, the biofabrication associated with the present invention seeks new technologies and the development of biomaterials that decrease the probability of loosening and reduce the wear of the prostheses.

Considering an injured joint surface, functional restoration of the joint can be facilitated if biomaterials that exhibit tribological behavior similar to those of natural joints are used when making prosthetic joint surfaces or grafting small defects. A class of materials that have interesting characteristics to mimic such behavior are polymer hydrogels, since they have physical similarities with the natural articular cartilage, mainly in their ability to deform when compressed, exuding fluid contained within. In addition, studying the lubrication, friction and wear of natural and artificial joints has allowed us to increase our understanding of how natural joints work and why they fail. The replaced joints are subjected to cyclic mechanical loads during movement, which are transferred by the tissues, which leads to wear, causing difficulties in predicting the interfacial stability between tissues and implants.

In the field of materials, a growing field is development, formation and testing in vivo and in vitro with biomaterials for medical applications. Defined as any substance, natural or artificial, that can be used in the human body to replace diseased or damaged parts, biomaterials should be tissue compatible and not produce toxic substances. In addition, they must be biofunctional, that is, they must have the proper characteristics to fulfill the desired function for the desired time; be sterilizable and also biotolerant. In selecting the material to be used, consideration should be given to its physical, chemical and mechanical properties, especially wear and fatigue resistance, modulus of elasticity, surface roughness, permeation rate, biostability, water absorption and bioactivity (Ferreira, 2007).

Biomaterials can be ceramic, metallic and polymeric, or a combination thereof, and can be applied as devices for controlled drug delivery systems; cement; dental implants; orthopedic fixation; scaffoids; sutures; eye, knee, hip and tendon prostheses; heart valves; catheter as drain; hernia correction mesh; cell encapsulation, joint cartilage substituents, among others.

The application of polymers in the biomedical area began with the use of celluloid for surgical implantation in the repair of skull defects, followed by the application of bakelitis in hip arthroplasty (Rosiak and Ulanski, 1999). However, the development of these biomaterials did not take into consideration their purity and biostability, which caused adverse reactions in the organism. Today, these characteristics are essential for application of the material to the human body and numerous polymers such as polymethyl methacrylate (PMMA), polyvinyl alcohol (PVAI), poly 2-hydroxyethyl methacrylate (pHEMA), polyvinyl pyrrolidone (PVP) etc, come gaining prominence in the biomedical area, due to the overcoming of these limitations. In addition, ultra high molecular weight polyethylene (PEUAP) and polyurethane (PU) can be studied for orthopedics applications due to their mechanical properties such as high abrasion resistance and low coefficient of friction.

In the biomedical field, polymer hydrogels are the front line of recent research. The great interest of researchers for these biomaterials lies in their characteristics and similarities with the body's soft tissues. These materials have a soft and elastic consistency; good mechanical strength; biocompatibility in contact with blood and body fluids and allow for cell nutrition. Defined as three-dimensional, cross-linked chain polymers that have the ability to swell without dissolving, they are commonly used as substituents of the articular cartilage, by improving its mechanical properties and as a support for cell growth in tissue engineering. But they can also be studied for use as bandages; contact lenses; vascular grafts; hemodialysis membranes, among others.

The principle of obtaining polymeric hydrogels is not recent, since 1960 the biomedical potential of these materials has been considered. Witcherle (1971) studied poly (2-hydroxy ethyl methacrylate) - pHEMA for use in ophthalmology, in the development of contact lenses. From then on, numerous studies in different areas and with different contributions began to be studied, with these polymers as potential targets.

One of the most important conclusions from previous work was the observation that the optimal mechanical performance of poly 2-hydroxy ethyl methacrylate (pHEMA) hydrogels for application as articular cartilage is directly related to the mechanical resistance of the substrate to which , is being supported. When the hydrogel is implanted by covering mechanically resistant substrates, such as Ultra High Molecular Weight Polyethylene (PEUPAM), its mechanical behavior becomes ideal for the desired application, which is to support and distribute the applied load during the movement, favoring the formation of a hydrogel. lubrication regime between surfaces (Bavaresco et al., 2008 (a); Garrido, 2007).

Moreover, this research has shown that pHEMA hydrogels have good physical-mechanical properties when subjected to tribological tests of friction and wear on the physiological conditions of a natural joint, stating their potentiality for use in repairing small joint surface defects, where Tissue behavior around and within the graft was evaluated for long post-implantation periods (Bavaresco et al., 2008 (a); Bavaresco eti a /., 2008 (b); Bavaresco et al, 2004 (a) ; Batista et al., 2008; Garrido, 2007; Malmonge, 2002; Malmonge and Belangero, 2002).

Hydrogels may be obtained in different forms depending on the desired application. Can present themselves as movies optically transparent; spongy, non-spongy gels, among others (Lannuzzi et al., 2010; Eljarrat-Binstock et al., 2007). Moreover, they are easily synthesized by different techniques, such as: (i) copolymerization: normally, one monomer is hydrophobic and the other hydrophilic. In this case, dissolution of the network is prevented due to the presence of ionic bonds or hydrophobic interactions (Song et al., 201 1; Wang et al., 2010; Gholap et al., 2004; Barcellos et al., 2000). But the use of copolymerization usually includes the use of toxic solvents and / or monomers, which may interfere with the biocompatibility of the final products obtained; (ii) heat treatment (freeze-thaw): Intermolecular interactions probably lead to the formation of hydrogen bonds, forming entanglements between the chains and, consequently, the formation of three-dimensional lattice via formation of crystallites, which act as cross-links (Hu eí al. , 2010; Gupta et al., 2010; Ru-Yin and Dang-Sheng, 2008). This technique may hinder adhesion between the surfaces of polymeric materials. During coating of the hydrogel under the substrate, thermal expansion may occur on the most mechanically resistant material given freeze-thaw cycles. (iii) radiation: allows obtaining hydrogels in a single step, with simultaneous cross-linking and sterilization (Singh and Pai, 201 1; Sahiner et al., 2006; Ulanski et al., 2002; Lugao and Malmonge, 2001; Martens and Anseth, 2000), thus being a tool used in the present invention, by coupling the infrared laser in the biofabrication system.

The use of radiation technique for the synthesis of polymer hydrogels was initially developed by Charlesby (1960) and Chapiro (1962), and its use for biomedical applications has been extensively studied (Hill et al., 201 1; Zainuddin et al., 2011; Wang et al., 2011; Tomic et al., 2010; Zhao et al., 2010; El-Din and El-Naggar, 2005; Sahiner et al., 2006; Zhain et al., 2002; Bhattacharya, 2000). Considering that one of the basic requirements for hydrogels to be in contact with body fluids is the absence of toxicity, then the radiation technique becomes an advantageous alternative. In addition, it is an efficient sterilization tool, allowing that the process of obtaining the hydrogel is carried out in one step, with formation and sterilization at the same time. This allows for simplification of technology and reduction of production costs. In addition, the use of radiation involves other advantages, such as:

· Absence of chemical agents, allowing to obtain materials of high purity, without contamination by residues of crosslinking agent or chemical initiators, eliminating possible cytotoxicity;

• Process control. The processes of initiation and termination of crosslinking occur simply by introducing and removing material from the radiation source;

• Easy to modify physical and chemical properties of hydrogels. That is, the mechanical properties of hydrogels can be altered by simply modifying the type of radiation, by adjusting the intensity and / or exposure time of the material (radiation dose).

The adhesion between two polymers, for application in joint prostheses, is presented as the limiting factor to the development of soft-layered prostheses. Such prostheses are medical devices that have two polymeric materials attached together. One has high mechanical strength and the other is compliant and soft, with properties similar to natural articular cartilage. The purpose of adhesion between these materials is to enable a lubrication regime between the surfaces in contact, without tearing from one surface to another. This may be possible thanks to the compliant polymer (hydrogel), which, when subjected to a charge, releases fluid contained within, allowing for reduction of friction and wear between the joint surfaces via the formation of an elastrohydrodynamic action film similar to synovial fluid from natural joints.

Without the presence of this film, there is a wear process in these devices, during the movement cycles, causing deterioration of the bone-implant interface, consequent failure of the prostheses and increase in the rate of revision surgeries in the patients. Usually, polymeric coatings or hydrogel shrink when dry due to chemical and physical changes, while the substrate (high wear-resistant polymer) tends to be rigid so that its surface area does not change when in the presence of body fluids. This causes the adhesive and cohesive process of the system to be impaired and fracture may occur between the surfaces. To ensure more efficient coatings, the surfaces must be chemically bonded.

New engineering concepts that enable the coating of surfaces with desired geometry and alternative materials that have mechanical and physical characteristics similar to the region being replaced or repaired are being continually researched.

Bose and Lau, 201 1 used the vapor deposition technique to obtain poly 2-hydroxy ethyl methacrylate hydrogels using a solvent free polymerization procedure. But in this article, there is no description of obtaining hydrogels in specific or three-dimensional geometries. Moreover, it does not take into account the adhesion of two polymeric materials and mechanical properties of the obtained products.

Kubinová et al, 2009 describe a new technique for obtaining poly 2-hydroxy ethyl methacrylate cholesterol scaffolds for tissue engineering application, demonstrating that the product obtained showed improvements in bioactivity, proliferation and cell adhesion properties. But it does not mention the possibility of obtaining devices with specific geometries.

Wolf et al, 2009 describe an orthogonal copolymerization strategy for the preparation of amphiphilic copolymers using a bifunctional primer. However, in this article, there are no reports of obtaining 3D structures. In addition, the use of chemical initiators can generate toxic residues in the final product, which feature is minimized with the present invention which uses infrared laser for process control, simultaneous sterilization and polymerization. Bártolo et al, 2004 propose a new bioprototyping process for the production of three-dimensional alginate scaffolds and cell encapsulation. However, unlike the present invention, there are no reports of adhesion between two materials.

One of the most common technologies for obtaining polymeric biomaterials is 3D printers. As described by Lipson and Kurman (Lipson and Kurman, 2010) this type of technology uses additive methods, depositing the raw material layer by layer to obtain the final product systematically. The material (metal, ceramic or polymer) can be extruded through a syringe or laser sintered.

Due to layer-by-layer material deposition, 3D technology is capable of combining various materials and textures, which normally cannot be combined on conventional machines. When working with raw materials that are chemically incompatible or require different manufacturing conditions, traditional production machines must work with the materials in separate processes and then assemble them.

3D printers have a clean manufacturing process, do not involve cutting, burning or scraping the material, producing little manufacturing waste. Thus, due to its precision and versatility, this technology has been gaining importance in several industrial segments: prototyping, virtual modeling and even the medical field, aiming at obtaining devices to improve people's quality of life, by replacing diseased or diseased organs. damaged. It is a process of "co-fabrication," not unlike biological growth, where hard and soft tissues are "co-fabricated" and interconnected into living beings of infinite complexities.

Given the great versatility of 3D printers, studies have focused on their possible application in the medical field, aiming to achieve the complexity of geometric shapes of organs and tissues. Biomaterials should be able to mimic living structures both in function and in thus, it is possible to replace damaged tissues (Jardini et al., 2010).

3D technology creates a physical object from a digital file. Initially the object is copied (scanned) to obtain the three-dimensional surface of the structure via tomography, resonance or scanning. With the 3D surface obtained, there is the generation of a virtual model, which is sent to the 3D printer for physical reproduction. 3D data also makes it possible to build models that will guide the development of products that are suitable for the human body. Thus, it becomes a scientific advance, directed to bioprinting or biofabrication technologies, considering the use of biomaterials.

Based on similar physical principles as the 3D printer, the present invention aims to synthesize polymeric hydrogels with physicochemical characteristics similar to articular cartilage, in specific geometries, to cover substrates to improve mechanical adhesion of the system.

In the prior art, the possibility of using polymeric hydrogels as a compliant surface (cartilage) in artificial joints presents as limiting factor the adequate adhesion between the interface of the elastomeric layer (hydrogel) and the support interface (substrate) (Burgess et al ., 2008). Normally, polymeric coatings shrink when dry due to chemical and physical changes, while the substrate tends to be rigid so that its surface area does not change when swollen. This causes the adhesive and cohesive process of the system to be impaired and fracture may occur between the surfaces. To ensure more efficient coatings, surfaces must be chemically adhered, although the interfacial wear between the coating and the substrate is relatively high for incompressible materials (Matthewson, 1982). However, appropriate material modifications, design considerations and effective manufacturing techniques as proposed in this invention, improved adhesion between the substrate and the elastomeric layer, encouraging the clinical use of these prostheses (Jones et al., 2009).

WO201 1038373 of 31/03/201 1 (Three-dimensional bioprinting of biosynthetic cellulose (BC) implants and scaffolds for tissue engineering cross-reference to related application, Gatenholm Paul, Backdahl Henrik, Tzavaras Theodore Jon, Davalos Rafael, Sano Michael ) describes a system and method for the production of three-dimensional biomaterials employing natural polymers and fermentation techniques. However, there is no reference to adhesion between two polymeric materials, and, disadvantageously, does not use medical imaging and rapid prototyping data to control the thickness of the material. Already the present invention allows to obtain three-dimensional structures from medical imaging data, obtaining control of thickness and volume of the deposited material. It also enables polymerization and crosslinking of polymers to be achieved through the use of infrared laser with defined geometries.

Another patent application which also describes a three-dimensional structure printing system is US2010278952 of 4/01/2010 (Silverbrook Kia). This type of technique does not consider materials such as gels or solutions. The high viscosity of the material to obtain the object, around 10 cP is a major problem encountered, ie, for object printing it is necessary that the material is at least prepolymerized. The present invention differs from the US patent application in that it uses low viscosity solutions (1.07g / cm ³ ) such as 2-hydroxyethyl methacrylate (HEMA) solutions, allowing polymerization prior to curing of the material. The material is deposited in a volume defined by a deposition syringe and the reaction occurs with the use of infrared laser, acting as a heat source. Thus, two polymers are adhered without necessarily both being prepolymerized.

In US201 10177590 of July 21, 2001 1

(Bioprinted Nanoparticles and Methods of Use, Clyne Alisa Morss, Buyukhatipoglu Kivilcim, Chang Robert, Sun Wei) a three-dimensional cell complex is biofabricated by deposition. There are no reports of adhesion between two polymeric materials and use of infrared laser as a heat source. Application of infrared laser in the polymerization and restricted curing of polymeric materials and interfacial adhesion between them is the differential of the present invention.

WO2007124481 of 01/01/2007 (Bioprinting three-dimensional structures onto microscale tissue analogues for pharmacokinetic study and other uses, Sun Wei, Chang Robert, Starly Binil, Nam Jae) describes a microfluidic system for monitoring and detecting changes in an input parameter of a substance, which includes a microfluidic device having a tissue chamber and a substitute tissue in that same chamber. Specifically, the invention relates to an in vitro model for studying pharmacokinetics and pharmaceutical applications, among other uses. This invention is subjected to a microfluidic device to mimic fluid conditions of a mammalian body, including the use of cells, which differs from the present invention.

WO201 1 107599 of 09/09/201 1 (Bioprinting station, assembly including such bioprinting station and bioprinting method, Guillemot Fabien, Catros Sylvain, Keriquel Virginie, Fricain Jean-Christophe) reports on a Bioprinting adapted to deposit standard biological materials, including cells, biomaterials, nanoparticles, drugs and more. In this case, the laser is coupled to transfer the biological material to a determined area under the substrate. The present invention utilizes infrared laser as an energy source for polymerization and curing between two polymeric materials.

Therefore, in view of the described technologies, the great advantage of the present invention is that it is capable of depositing polymeric hydrogels layer by layer, having the infrared laser as the heat source, which is responsible for initiating the polymerization and crosslinking reactions. This technology also allows the coating to be carried out under geometries which may be obtained via computed tomography (CT), scanning or resonance (MTI). The data is transformed into a standard STL file (physical model that approximates the surface of the solid in triangular format), allowing the system to be controlled in the desired geometry.

The use of infrared laser, or better, the coupling of a laser fiber in the mechanical XYZ axis scanning system has the advantages of being flexible and allowing the control of the temperature and location of the radiation, allowing a better incidence of the beam over the sample. This condition enables the energy deposited in the hydrogel to have the function of crosslinking and reaching the substrate in order to improve adhesion between surfaces.

In view of the foregoing, the present invention has advantages in several respects. Firstly, it proposes a new process for obtaining polymeric biomaterials employing an infrared laser that ensures fast, restricted and localized cure. In addition, this process allows the deposition of layer-by-layer hydrogels and, once the energy flow deposited in the sample is controlled, the polymeric curing occurs in a defined volume, allowing the obtaining of specific geometries. Another crucial point for the process is the use of infrared leisure in the XYZ axis mechanical scanning system which is flexible, allowing fine temperature control, avoiding heat dissipation in undesired regions and consequently cross-linking in the adjacent regions. This enables you to control the thickness of the overlay and reduce process costs by minimizing system losses.

Thus, with the process described herein it is possible to use the laser system to thermally cure other types of polymers, since infrared laser irradiation generates heat which is the driving force for initiating and propagating thermal curing. In addition, the laser enables healing localized once the laser energy is confined to the diameter of the laser beam.

The system also allows the design of the substrate in any geometry or anatomy of the human body, ie, from digital medical data provided by tomography or magnetic resonance, a software performs the treatment of medical images and generates a file of the structure in 3D. allowing the construction of the body part exactly in the dimensions of the desired human part.

Highlighting the products obtained in the present invention, such polymeric biomaterials have a good adhesion between the polymeric hydrogel interfaces and the substrate, mimic the behavior of a natural joint, reducing the friction and, consequently, the wear and improvement of the physicochemical properties. chemical materials. The product obtained can have free geometry and can be constructed with desired geometry even with internal cavities and controlled porosity, a fact that is impossible to obtain in conventional machining and material forming processes.

Among the various applications that these biomaterials present, one of them is the use in the manufacture of biological substitutes and medical devices. In addition, they may be employed as polymeric hydrogels for cell support and growth, scaffolds and polymeric coatings for use in artificial joint prostheses.

Brief Description of the Invention

The present invention relates to a process of obtaining polymeric biomaterials by the adhesion of two polymers employing infrared laser.

The invention describes a process comprising the steps of medical data acquisition, fabrication of the physical model in 3D geometry, substrate preparation using machining, followed by a coating step comprising the addition of reagents, deposition of the polymeric hydrogel layer by layer and finally the polymerization steps and crosslinking simultaneously. Furthermore, the present invention relates to the use of said biomaterials.

Brief Description of the Figures

The structure and operation of the present invention, together with further advantages thereof may be better understood by reference to the appendices and the following description:

- Figure 1 shows the machining of the substrate, where (a) is the sketch, (b) the PEUAPM substrate in flat geometry and (c) is the PEUAPM substrate in cylindrical geometry obtained after machining.

- Figure 2 shows the PEEMAM substrates covered with pHEMA hydrogels, where (a) is the blistered surface and (b) the nonuniform thickness.

- Figure 3 shows the flat geometry PEUPAM substrates covered with pHEMA hydrogels at different concentrations of the HEMA monomer, where (a) solution X, (b) solution Y, (c) solution Z

- Figure 4 shows the flat geometry PEUAPM substrates coated with pHEMA hydrogels at different concentrations of the crosslinking agent, diethylene glycol dimethacrylate (DEGDMA), where (a) solution A, (b) solution B and (c) solution C .

- Figure 5 shows the cylindrical geometry PEUAPM substrates covered with pHEMA hydrogels (solution A).

- Figure 6 shows a flat geometry polyurethane substrate covered with pHEMA hydrogel (solution Z).

- Figure 7 shows the micrographs of the PEUAPM coating with pHEMA hydrogel overlay at different concentrations of HEMA, with 500x magnification, where (a) is solution X and (b) is solution Z.

- Figure 8 shows the micrographs of the interface of PEUAPM coating with pHEMA hydrogel at different concentrations of cross-linking agent, DEGDMA, with 500x magnification, where (a) is solution A, (b) is solution B and ( c) is solution C. Detailed Description of the Invention

The present invention describes a process of obtaining polymeric biomaterials by coating a substrate by a polymeric hydrogel employing infrared laser.

Biomaterials obtained by the process described in this invention comprise layer by layer deposition of polymeric hydrogel under a substrate.

An object of the present invention is a process for obtaining the products described above comprising the following steps:

a) Acquisition of medical data

b) Fabrication of the physical model in 3D geometry c) Substrate preparation

d) Machining

d) Coating

d1) Addition of reagents

d2) Deposition

d3) Polymerization

d4) Crosslinking

The first step (a) of medical data acquisition is done using technological equipment in the medical field, such as X-rays, computed tomography (CT) and magnetic resonance imaging (MRI), which allow to obtain internal images of the human body. These tools are commonly used to visualize bone, organ, and tissue configurations and provide additional information on medical imaging in electronic format (DICOM). Through these electronic files (DICOM), the second step (b) of the process is performed. Physical models of the structure of the human body are obtained via rapid prototyping technique. The data is digitized and converted, sliced, to standard STL file (physical model that approximates the surface of the solid in triangular format). Then the model is evaluated and validated. In the third step (c), the substrate preparation is performed using the machining (d) in specific geometry to be covered. For the preparation of the substrate in step (c), Ultra High Molecular Weight Polyethylene (PEUAPM) with molecular weight of 2.5 million g / mol and density of 0.6 g / cm ³ and Polyurethane with different degrees of polydispersity, from different chemical routes. The substrate is machined (d) in flat plates 36 x 32 x 3 mm, containing two holes with a diameter of 1 to 8 mm, preferably 5 mm, at a distance of 10 to 30 mm, preferably 22 mm. or in cylindrical geometry 4 mm in diameter as shown in Figure 1.

Then, in the fourth step (d), the chemical reagents for polymerization and cross-linking of the hydrogel under the substrate are added to a syringe with pump-adjustable rate and deposition volume for formation of the polymer hydrogel. The hydrogel is prepared by the addition of the following reagents (d1): a monomer or polymer, belonging to the group of lactones, alcohol or methacrylates, preferably 2-hydroxy ethyl methacrylate (HEMA), in a range of 20 to 100% w / w preferably 80% w / w; a crosslinking agent selected from diethylene glycol dimethacrylate (DEGDMA), trimethylolpropane trimethacrylate (TMPTMMA), Ν, methylene bis bis acrylamide, ethylene glycol dimethacrylate (EGDMA), triethylene glycol dimethacrylate (TEGDMA), and other cross-linking agents di, tri and tetra functional, preferably diethylene glycol dimethacrylate (DEGDMA), between 1 and 3% w / w with respect to the HEMA monomer content, preferably 1% w / w; thermoinitiators, selected from 2,2'-Azobis (2-methylpropionitrile) (AIBN), dibenzoyl peroxide and potassium persulphate, preferably potassium persulphate (PKS) and dibenzoyl peroxide, preferably a 1% w / w. All reagents are physically mixed and kept under agitation for complete homogenization.

Then the hydrogel obtained in step (d1) is deposited (d2) layer by layer under the substrate, in specific geometry delimited by 3D CAD model, employing a prototyping equipment coupled to it an infrared fiber optic laser. The laser acts as a heat source and is responsible for initiating the hydrogel polymerization and cross-linking reactions under the substrate, allowing for coating and adhesion between two polymeric surfaces. That is, the laser performs the function of ensuring polymerization and crosslinking, in addition to mechanical embracing in the anterior layer. From 1 to 10 layers of polymeric hydrogel, preferably 3, are applied between 10 and 60 minutes, preferably 35 minutes, with solution flow rate between 10 and 120 mL / hr, preferably 110 mL / hr. The laser scanning speed ranges from 50 to 100 m / s, preferably 100 m / s and the power ranges from 10 to 40 W, preferably 30 W.

The polymeric hydrogel, in desired three-dimensional geometry, undergoes radical polymerization (d3) followed by crosslinking (d4) in a single step. Initially, linear molecules of the polymer are formed. With increased monomer conversion and the presence of the crosslinking agent, free radicals react with double bonds producing chemical crosslinks between unconnected polymer chains, leading to hydrogel formation in specific geometry. Therefore, hydrogel synthesis under the substrate is obtained in a localized manner, with control of the laser intensity, power, scanning speed, volume and deposition height of the solution. Porous hydrogels are obtained with reaction time ranging from 1 to 3 minutes, preferably 3 minutes and dense hydrogels within a range of 1 to 3 minutes, preferably 2 minutes. The wavelength of the infrared laser source ranges from 1000 to 2000 nm, preferably 1070 nm and the diameter ranges from 0.5 to 1.0 cm, preferably 0.8 cm. The laser power under the solution is kept constant between 29.5 and 30.5 W, preferably 30 W and the distance from the laser focus to the center point of the solution is estimated between 4.5 and 10.0 cm, preferably 9 W. .5 cm tall.

The polymeric biomaterials obtained at the end of the process described in this invention have biomedical potential and different physicochemical properties and can be applied as biological devices, support for cell growth in tissue engineering, due to biocompatibility and non-toxicity of biomaterials, respectively, and as covering of rigid substrates, mimicking the characteristics of the natural articular cartilage.

Example 1: Coating PEUAPM Substrate with pHEMA Hydrogel

Using the process described in the present invention, commercial PEUAPM porous plates were coated with pHEMA hydrogels obtained from the concentrations described below:

• 40, 60 and 80% w / w monomer (2-hydroxy ethyl methacrylate); 1% w / w potassium persulfate (thermoinitiator) and 2% w / w diethylene glycol dimethacrylate (crosslinking agent) - Solution X, Y and Z.

• 100% monomer; 1% w / w dibenzoyl peroxide (thermoinitiator); 1, 2 and 3% w / w diethylene glycol dimethacrylate (DEGDMA, cross-linking agent) - Solution A, B and C.

Process variables such as volumetric flow of HEMA solution (flow) and sweep speed were analyzed. The height of the syringe relative to the substrate and the infrared laser to the substrate were kept constant at 4.5 cm and 9.5 cm, respectively. Table 1 shows the evaluated operating conditions.

Table 1 Operating conditions assessed at the overlay stage

Variables Value (s)

Total volume of syringe 60 ml

Syringe Diameter 29.45mm

Rate 10 to 120 ml / h

Sweep Speed 50 and 100 m / s

Another important aspect analyzed, for the two concentrations mentioned, concerns the sweep rate and flow rate of the hydrogel. Lower scanning speeds, 50 m / s, at flows between 10 and 50 mL / h, increased laser residence time under the substrate, leading to solution evaporation and, consequently, PEUAPM fusion. At flow rates above 50 mL / h the evaporating effect of the solvent was bypassed. However, there was a larger amount of solution being deposited at the same time of laser incidence (scan). This caused occasional overheating of the solution, leading to blistering, uneven surface coatings and uneven thicknesses. In addition, it was found that at low sweep speeds, pHEMA polymerization and crosslinking at the syringe nozzle was a more pronounced effect, leading to clogging of the system.

For sweep speeds of 100 m / s, HEMA solution flow rates between 10 and 100 ml / h provided discontinuous flows. The deposition of the material by the syringe did not continually follow the path taken by the laser. Surface points were heated without the presence of the solution. This led to spot fusion of PEUAPM substrate. Under this condition, the substrate temperature was randomly measured by another infrared laser source. Results demonstrated point temperatures of 165 ° C, above the melting temperature of PEUAPM.

Continuous flows, however, were observed at flows of 1 10 and

120 mL / h for this same sweep rate. But at high flow rates (120 mL / h), there was an excess of solution under the substrate, with significant material loss and higher energy required to obtain the hydrogel.

The total volume of solution inside the syringe also showed considerable effect. At high sweep speeds, there was a shortage of solution before the deposited layer was finished. The lack of solution led to air formation in the syringe and consequent discontinuous flows. To minimize such effects, a flow rate of 10 mL / h and a sweep speed of 100m / s were selected, with solution deposition intervals. Still, the thickness of the crosslinked layer (hydrogel) and the presence of bubbles were not controlled parameters, as shown in Figure 2.

PHEMA hydrogels under PEUAPM were then obtained via 3-step system scan. Steps 1 and 3 involved the deposition of the material (HEMA solution) and heating via infrared laser source continuously. Step 2 consisted only of passing the laser over the substrate without the addition of HEMA solution. The laser had the function of ensuring polymerization and crosslinking, as well as mechanical embracing in the anterior layer. Total process time was 35 minutes. Figures 3 to 6 show pHEMA hydrogel coated PEUAPM substrates using the process described in the present invention.

For both concentrations evaluated, pHEMA hydrogel was found inside the PEUAPM pores. That is, hydrogel formation not only on the surface but also internally on the substrate, although hydrogels with 1% w / w DEGDMA (solution A) showed better adhesion. Figures 7 and 8 show the hydrogel-PEUAPM interface, after coating for all studied concentrations, at 500x magnification.

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Claims

1 . Process for obtaining polymeric biomaterials comprising the following steps:

a) Acquisition of medical data

b) Fabrication of the physical model in 3D geometry c) Substrate preparation

d) Machining

d) Coating

d1) Addition of reagents

d2) Deposition

d3) Polymerization

d4) Crosslinking

Method according to claim 1, characterized in that step (a) comprises the acquisition of medical data by X-ray, computed tomography or magnetic resonance.

Process according to Claim 1, characterized in that step (b) is preferably carried out via the rapid prototyping technique.

Process according to Claim 1, characterized in that step (c) comprises a substrate selected from an ultra high molecular weight polyethylene or a polyurethane with different degrees of polydispersity.

Process according to Claim 4, characterized in that the ultra high molecular weight polyethylene preferably has a molecular mass of 2.5 million g / mol and a density of 0.6 g / cm ³ .

Process according to claim 1, characterized in that step (d) comprises machining flat substrate plates of dimensions 36 x 32 x 3 mm, containing two holes with a diameter of 1 to 8 mm, preferably 5 mm, at a distance between them from 10 to 30 mm, preferably 22 mm or machining substrates with 4 mm diameter cylindrical geometries.

Process according to Claim 1, characterized in that step (d1) comprises the addition of a monomer or polymer, a cross-linking agent and a thermo-initiator.

Process according to Claim 7, characterized in that the monomer or polymer is selected from the group of lactones, alcohol or methacrylates, preferably 2-hydroxyethyl methacrylate.

Process according to Claim 8, characterized in that the 2-hydroxy ethyl methacrylate is added in a range of from 20 to 100% w / w, preferably 80% w / w.

Process according to Claim 7, characterized in that the cross-linking agent is selected from diethylene glycol dimethacrylate, trimethylolpropane trimethacrylate, Ν, ileno 'methylene bis bis acrylamide, ethylene glycol dimethacrylate, triethylene glycol dimethacrylate, among others. di, tri and tetra functional chemical agents, preferably diethylene glycol dimethacrylate.

Process according to Claim 10, characterized in that the diethylene glycol dimethacrylate is added between 1 and 3% w / w in relation to the monomer or polymer content, preferably 1% w / w.

Process according to Claim 7, characterized in that the thermoinitiator is selected from 2,2'-Azobis (2-methylpropionitrile), dibenzoyl peroxide and potassium persulphate, preferably potassium persulphate and dibenzoyl peroxide.

Process according to Claim 12, characterized in that the thermoinitiator is added between 1 and 3% w / w, preferably at 1% w / w.

Process according to Claim 7, characterized in that said reagents are physically mixed and kept under agitation for complete homogenization.

Process according to claim 1, characterized in that step (d2) comprises depositing 1 to 10 layers of polymeric hydrogel obtained in step (d1), preferably 3, between 10 and 60 minutes, preferably 35 minutes, with solution flow rate between 10 and 120 mL / hr, preferably 110 mL / hr, with laser scanning speed ranging from 50 to 100 m / s, preferably 100 m / s and power ranging from 10 to 40 W, preferably 30 W.

Process according to claim 1, characterized in that steps (d3) and (d4) occur simultaneously.

Process according to Claim 16, characterized in that the final time of the porous hydrogel reactions ranges from 1 to 3 minutes, preferably 3 minutes.

Process according to Claim 16, characterized in that the final time of dense hydrogel reactions ranges from 1 to 3 minutes, preferably 2 minutes.

Process according to Claim 16, characterized in that the wavelength of the infrared laser source ranges from 1000 to 2000 nm, preferably 1070 nm, the diameter ranges from 0.5 to 1.0 cm, preferably 0.8 cm. cm the laser power is kept constant between 29.5 and 30.5 W, preferably 30 W and the distance from the laser focus to the center point of the solution is between 4.5 and 10.0 cm, preferably 9.5 cm high.

Polymeric biomaterials characterized in that it is obtained according to the steps described in claims 1 to 19.

21 Use of the polymeric biomaterials described in claim 20 characterized in that they are employed as support and growth of polymeric cells, scaffolds and coatings for use in artificial joint prostheses.