HK1230108B - Method for manufacturing a porous metal material for biomedical applications and material obtained by said method - Google Patents

Description

Method for producing porous metal materials for biomedical applications and materials obtained by said method

Technical Field

The present invention relates to porous metallic materials for use in certain biomedical applications. In particular, the present invention comprises methods for producing and alloying titanium having enhanced bone bonding and adsorption properties.

Background Art

A major issue in selecting metallic materials for biomedical arthroplasty applications is the need to combine functional characteristics such as osseointegration and stiffness similar to human bone, all in a material that can withstand dynamic loads in use.

Titanium has three main benefits: its excellent biocompatibility, reduced modulus of rigidity (110 GPa compared to 210 GPa for conventional sanitary steel), and compatibility with diagnostic and assessment technologies such as CAT scans or MRIs. All of this makes titanium or its alloys the most suitable metal material for manufacturing any prosthesis or implant that must be placed in the human body.

When forming the metal material used to manufacture these prostheses, using wrought materials that are machined into shapes designed by experts, the aforementioned mechanical properties are maintained.

However, even if these properties are superior to those of any steel or CrCo material on the market, they are still insufficient to improve two fundamental aspects, namely bone adsorption and bone integration. The first of these aspects is closely related to the difference between the rigidity modulus of bone (0.5 at 30 GPa) and the rigidity modulus of the metal prosthesis. As these two values approach each other, the prosthesis functions as bone and bone resorption decreases, increasing the life of the implant and thus improving the patient's quality of life. The simplest way to reduce the modulus without modifying the material is to increase the porosity of the system, and the most prominent technology for producing porous materials is powder metallurgy.

A second area where existing materials for prostheses could potentially improve is osseointegration. Uncemented prostheses are increasingly being used to reduce the impact of the anchoring itself, not only in porous cavities but also in assessing the risk of fracture of the set cement. The use of uncemented prostheses involves the development of the anchoring system itself, as well as the application of biocompatible coatings for ceramic or metallic materials. The most widely used today are microsphere deposition techniques using thermal spraying or by small-scale gluing and sintering. Both types of systems have several drawbacks. Thermal spraying produces a rough surface, but this surface roughness lacks substantial and deep interconnectivity, so the bone tissue merely "grasps" into the cavities and peaks of the resulting topography. Conversely, given the small number of weld points present in the microspheres to be sintered, microsphere deposition carries the inherent risk of some of these microspheres becoming dislodged, with consequent risks to the patient's health. Furthermore, in the latter case, there have been numerous issues with prostheses fracturing during use due to fatigue. The welded attachment of the spheres to the mold creates sharp edges, which serve as points of increased stress.

Processing titanium materials and alloys by powder metallurgy to obtain porous materials is known in the prior art. The greatest success has been achieved by eliminating the use of spacing agents in some subsequent production processes and seems to be on the verge of industrialization.

The use of a spacer limits the size of the pores that result in the component after the spacer is removed. However, the presence of macropores does not necessarily imply the presence of channels of similar size to the macropores for capillary growth toward the interior of the porous material.

The powder metallurgy process allows titanium and/or titanium alloy powders with an average particle size of 300 microns to be compacted to a size of less than 25 microns. This process, followed by appropriate sintering, enables the resulting titanium material to achieve a density between 85% and 98% of that of a solid body, due to the significant shrinkage observed during sintering. Given the fine grain structure, the mechanical properties, when the component is sintered under high vacuum conditions, are approximately the same as those of the solid material.

Stress is transmitted to the bone through the prosthesis, so the prosthesis must have an elastic modulus as similar as possible to that of the bone to properly transfer loads to the bone and prevent so-called mechanical stress shielding, which causes bone resorption due to the lack of stress applied to the bone. Therefore, if the prosthesis as a whole does not have a modulus similar to that of the bone, the surface elastic modulus is less important. If fine powders are used, the load transfer during uniaxial compression of the metal powder produces a density gradient that causes deformation of the initial pressed shape during high-vacuum sintering. This is a significant disadvantage because it makes machining to the final shape after the sintering step impossible due to the closure of surface voids and the obstruction of bone integration. The presence of deformation during sintering, and in particular the presence of a "skin" of laminated powder on the side surfaces of the component (due to friction during removal from the mold), has led to the decision to perform green processing to open the pores and find the most suitable dimensions for achieving the final shape and size after sintering. A certain material strength is essential for green processing, and some references and patents (e.g., US Pat. No. 7,674,426 B2) describe the use of cold isostatic pressing for this purpose. All of this makes the system more expensive and, furthermore, since these methods are performed with fine titanium or titanium alloy powders, the final interconnections between the pores do not exceed 10 microns, so vascularization and efficient bone growth in the porous system remain problematic.

In some embodiments, the present invention relates to a kind of implant that can be used for the treatment of a variety of diseases. For example, the implant of a kind of implant that can be used for the treatment of a variety of diseases is ...

Summary of the Invention

To address the aforementioned issues, the method of the present invention proposes the use of titanium powder with specific properties and its mixture with a salt of specific size in specific proportions. More specifically, the present invention proposes a method for obtaining a porous titanium component, characterized in that the initial titanium powder is pure, has an average particle size of 200 microns, a flow rate of 93 seconds, and an apparent density of 1.0 g/ ^cm³ , and is mixed with at least 50% by weight of NaCl having a particle size of 300 to 600 microns at a ratio of 34% by weight of titanium.

These features result in pore interconnections greater than 150 microns, and the biocompatibility of titanium-based materials, which makes the product well-suited for improving the material's bone integration while maintaining adequate fatigue strength.

BRIEF DESCRIPTION OF THE DRAWINGS

For the purpose of aiding a better understanding of the features of the present invention according to a preferred practical embodiment of the invention, a set of drawings are appended to the following specification, wherein the illustrative features of the following drawings are described.

FIG. 1 a is an electron microscopic image of irregular titanium powder having an average particle size of 200 μm used in the present invention.

FIG1 b is an image of the microstructure of a component manufactured according to the present invention, in which it can be seen that the grain size and the interconnection between the pores are greater than 150 microns.

FIG. 2 a is a graph showing the compression fracture behavior of a component manufactured according to the method of the present invention.

FIG. 2 b is a graph showing fatigue behavior of an intervertebral fusion cage manufactured according to the method of the present invention under use conditions.

FIG. 3 depicts the logarithm of the size of the pores relative to the average volume that can be occupied (invaded) by material for a structure manufactured according to the claimed method.

FIG. 4 is a graph illustrating the high crystallinity of parts produced according to the present invention.

FIG5 is a diagram showing the morphology of osteoblasts on the surface of a porous material.

DETAILED DESCRIPTION

The general aspects of the invention start with pure titanium powder having a particle size distribution of 45 to 300 microns, with more than 90% of the particles being between 75 and 250 microns and an average particle size of 200 microns. The powder has a flow rate of 93 seconds, an apparent density of 1.0 g/ ^cm3 , and a tap density of 1.25 g/ ^cm3 .

The flow rate of the material was calculated according to ISO 4490, the apparent density was calculated according to ISO 3923/1, and the tap density was calculated according to ISO 3953. The apparent density of 22% of the solid (dense) material density (1.0 g/ ^cm3 versus 4.51 g/ ^cm3 of solid titanium) indicates the presence of a highly irregular powder surface, which, together with the average particle size of 200 microns and the salt particle size of 300 to 600 microns, makes it feasible to achieve pore interconnect sizes greater than 150 microns using the pressing and sintering process.

Titanium is mixed at a rate of 34% by weight with 300 to 600 microns and 50 to 80% by weight of NaCl, with a binder added at a rate of at least 15 to 100%. The material is then subjected to a heat treatment to remove the binder and salt, followed by continuous rinsing in bidistilled water, and after pressing the material at 200 to 400 MPa, preferably 300 MPa, it is sintered at a temperature of 1200 to 1400°C (preferably 1300°C) and a pressure of less than 4×10 ⁻⁴ mbar. These parameters achieve not only suitable strength and homogeneity, but also an appropriate porosity and pore length.

As seen in Figure 1a, the titanium powder has an irregular morphology and a particle size greater than 150 μm.

As can be seen in Figure 1b, the structures of the developed material are fully interconnected with an interconnect size larger than 150 microns.

Figure 3 shows the plotted results of mercury porosimetry. In porous interconnected structures like this, the most frequent maximum intensity value corresponds to the size of the interconnected channels between the pores. Therefore, the strongest second peak represents the size of the interconnected pores, which are larger than 150 microns. The first peak represents all internal pores smaller than 10 microns, which have a specific biological function within the material and during the osseointegration process: serving as a "food" reservoir for cell growth.

To transform the titanium surface into a bioactive surface, the inert titanium dioxide (TiO ₂ ) layer on the implant surface (which spontaneously forms on titanium and its alloys) is reacted with a 5M alkaline solution of sodium hydroxide (NaOH). During the NaOH treatment, the TiO ₂ partially dissolves to form an alkaline solution due to the corrosive action of the solution's hydroxide (OH ⁻ ) radicals. Consequently, a sodium titanate (Na ₂ TiO ₃ ) gel layer forms on the surface. The alkaline reaction is then neutralized with H ₂ O at 60°C for 24 hours. This is followed by a heat treatment at 600°C for 1 hour to dehydrate, densify, and increase the substrate adhesion of the sodium titanate gel layer. This results in a stable, partially crystalline Na ₂ TiO ₃ layer that promotes bioactivity and improves surface properties.

Example Intervertebral Fusion Cage:

Pure titanium grade 2 powder with a particle size distribution between 45 and 300 microns and an average particle size of 200 microns was used. 65% by volume of NaCl with a size of 300 to 600 microns was introduced. 15% ethylene glycol was added to the final mixture and mixed for 10 minutes in a double-cone mixer. The wet mixture was introduced into a mold with the final geometry. Due to the uniform shrinkage of the material during sintering, the mold size was increased by 8%. Uniaxial pressing was carried out in a hydraulic press at 300 MPa, with excess binder used to promote uniform pressure distribution during demolding. This effect ensures a very uniform distribution of the pressing pressure, resulting in a uniform green density of the pressed material and no deformation in the final sintered part due to shrinkage differences. The pressed parts with titanium and salt were passed through an oven at 200°C for 6 hours to remove ethylene glycol residues (ethylene glycol evaporates at approximately 190°C). At 200°C, there is no risk of oxygen incorporating into the titanium structure, thus minimizing possible contamination by these elements. A subsequent bath cycle is performed to remove the spacer until the ionic conductivity stabilizes at a very low value or similar to that of distilled water used as a solvent. The cleaning is performed by applying a vacuum to accelerate the salt dissolution process. Once the salt has been removed from the component, it is properly treated and placed in an air oven at 120°C for 4 hours to completely dry it.

The samples are then sintered in a high vacuum furnace (<4×10 ⁻⁴ mbar) at 1200°C to 1400°C, preferably 1300°C, for 4 hours. Once sintered, the samples are machined to round the edges and reduce the likelihood of particle shedding. Finally, a cyclic washing process in distilled water, ethanol, and acetone is performed, all under ultrasound, to properly clean the parts.

Compression tests were carried out on treated and untreated samples, and no significant differences were observed in the results in terms of elastic limit, breaking load or elongation at break. However, a specific trend towards an increase in the elastic modulus of the material was observed when carrying out bioactive heat treatment. Mechanical compressive strength and elastic modulus (10 GPa) values and good fatigue behavior thereof can ensure that this type of porous material is used for the good mechanical behavior of the intervertebral fusion cage application of the spine, and demonstrates the possibility of many other applications for hard tissue replacement. Fig. 2 a and Fig. 2 b illustrate good monotonic bending and compression behavior under fatigue, simulating the work of the spine. In these mechanical tests, the infinite life value exceeds 350 kg during the service life and there is no particle shedding in any case.

To evaluate the ability of implant materials to form an apatite layer on their surface, in vitro testing was performed using samples produced according to the method of the present invention, following the guidelines of the international ISO 23317 standard (Implants for surgery – In vitro evaluation of apatite – Formation ability of implant materials). This test consists of immersing the material in a solution with an ion concentration, pH, and temperature nearly identical to that of blood plasma, known as simulated body fluid or SBF.

The bioactivity test with SBF is evaluated based on the formation of apatite on the surface due to ion exchange between SBF and the chemically and thermally treated surfaces.

When thermo-chemically treated titanium is immersed in SBF, the Na ⁺ _ions of _Na2TiO3 exchange with the _H3O ⁺ of the aqueous medium, and Ti-OH groups form on the metal surface. The Ti-OH groups formed on the surface combine with the calcium Ca2 ⁺ ions of the SBF to form amorphous calcium titanate ( _CaTiO3 ). The Ti-OH groups on the Ti surface are charged and cause Ca2 ⁺ ions to precipitate on the surface and bind to them.

Some calcium ions react with HPO ₄ ^2- phosphate in SBF to form calcium phosphate (CaP) on the implant surface. The release of sodium Na ⁺ ions along with H ₃ O ⁺ ions in SBF causes the pH of the solution to increase. This in turn produces greater CaP ion activity, leading to rapid deposition of apatite on the titanium surface.

After the thermo-chemical treatment, the sample exhibited a microstructure such as that observed in Figure 4. Since the treatment was liquid, it penetrated the entire implant surface, resulting in the sodium titanate covering the entire surface, even into the pores. Because it was not a coating process, but rather a titanium-based crystallization process, a high degree of crystallinity was observed, ensuring adhesion to the implant. This prevented possible bacterial penetration, which occurs in implants with bioactive coatings, as several micrometers exist between the layer and the substrate, allowing bacteria and microorganisms to colonize.

Samples treated with different amounts of alkaline solution and subsequently by heat treatment were immersed in SBF for 10 days to determine surface bioactivity, as shown in the ISO 23317 standard. The crystal structure of the apatite formed can be confirmed after immersion. It is very important that apatite is crystalline. In implants coated with hydroxyapatite by plasma, the apatite is mostly amorphous. This causes the apatite layer to be dissolved very quickly by the physiological medium, while the implant and bone are kept apart at a certain distance so that osseointegration does not occur. In this case, the apatite is fully crystallized and its dissolution is much slower than amorphous apatite, which is optimal for the osseointegration process.

Based on the absorbance values obtained after incubation of SAOS-2 osteoblasts for 24 hours and 10 minutes with LDH reagent, it can be concluded that:

No cytotoxic effects related to the indirect exposure of SAOS-2 cells to the 5 concentrations of extracts obtained from the analyzed samples were observed.

In all samples, the absorbance values relative to the negative control were higher than 75%, and thus the values were within the allowable range for assuming that cytotoxicity was absent.

Between the 1st day and the 14th day, an increase in proliferation was observed because the cells had not yet begun to differentiate and proliferate. A certain decrease in proliferation was subsequently observed, which was (almost absolutely certain) attributed to the increase in cell differentiation after the 14-day incubation. All of these behaviors are normal in such cells. ALP activity also increased after 7 days, indicating cell differentiation. ALP is an indicator of the beginning of differentiation, and it is believed that after the increase in phosphatase activity after 7 days of incubation, it is normal for phosphatase activity to decline after 14 days of cultivation. This phenomenon is highly classified as a typical early cell differentiation process in the scientific literature. As the culture time increases, the number of cells and the degree to which they penetrate into the material also increase. The image of the sample hatched for 14 days has shown the degree of complete penetration into the sample, which is equivalent to half of the thickness of the component.

Figure 5 shows the morphology of osteoblasts from AMES on the surface of the porous material with a very high degree of focal points, which ensures good cell attachment and health, as shown by the osteocalcin levels found. The attachment, proliferation and differentiation in contrast to the high osteocalcin and gene expression levels ensure accelerated bone integration.

Claims (7)

1. A method for obtaining a porous titanium component, comprising: Providing pure initial titanium powder having a particle size of 45 μm to 300 μm, a flow rate of 93 seconds calculated according to ISO 4490 standard, and an apparent density of 1.0 g/cm ³ calculated according to ISO 3923/1; mixing the initial titanium powder in a ratio of 34 wt % with at least 50 wt % of sodium chloride having a particle size of 300 to 600 microns to obtain a mixture; and Obtaining a biocompatible porous titanium component for biomedical applications by sintering the obtained mixture at a temperature of 1200° C. to 1400° C., having a pore interconnection size greater than 150 microns, The method includes the steps of adding a binder, pressing the material, and washing. 2. The method according to claim 1, further comprising: i) adding a binder in a proportion of at least 15% by weight to said mixture comprising said initial titanium powder and sodium chloride; ii) pressing the resulting material obtained in step i) at 200 MPa to 400 MPa; iii) removing the binder and sodium chloride by heat treatment followed by continuous washing of the resulting mixture in bidistilled water; iv) sintering the resulting mixture obtained in step iii) at said temperature of 1200° C. to 1400° C. and a pressure of less than 4×10 ⁻⁴ mbar; v) Machining the surface of porous titanium parts. 3 . The method according to claim 2 , wherein the porous titanium component is introduced into a 5M sodium hydroxide (NaOH) alkaline solution after step v). 4. The method according to claim 2, wherein the binder is ethylene glycol. 5. The method according to claim 2, characterized in that step i) is carried out in a double-cone mixer, and the flushing of the mixture in step iii) is carried out by uniaxial die pressing. 6. The method according to claim 5, characterized in that the uniaxial die pressing is carried out in a hydraulic press. 7. The porous metal material produced according to claim 1.