JPH025425B2 - - Google Patents

Description

[Detailed description of the invention]

TECHNICAL FIELD The present invention generally relates to improved prosthetic devices. In one aspect, the invention relates to a prosthetic device with a sintered thermoplastic coating. In another aspect, the present invention is directed to an improved prosthetic device in which a porosity or density gradient exists in the sintered thermoplastic coating. Furthermore, the present invention is directed to a method of manufacturing such a sintered coating. Prior to the present invention, various methods for attaching prosthetic devices to the musculoskeletal system have been disclosed in the literature.
These methods can be classified as (1) impaction, (2) nails and screws, (3) cement, and (4) porous surface materials. Although the use of porous surface implants for fixation purposes is recognized as potentially offering significant advantages, this technique is not generally accepted in the surgical field. This is because issues of early fixation and long-term stability are relevant for this type of device. The various devices known to date include:
No. 3,986,212, issued May 19, 2006, which describes a composite prosthetic device containing a porous polymeric coating to obtain bone fixation by tissue ingrowth. ing. Porous polymeric materials that have been shown to be useful are those having interconnected pores of a specified density and average pore size.
The disclosed polymeric materials include high density polyethylene and polypropylene or mixtures thereof with certain critical parameters. It has also been shown that the coating can be mechanically interlocked or chemically bonded to the device. U.S. Patent No. issued July 27, 1976
No. 3971134 relates to a fixed denture for permanent or long-term implantation into the mandible of a living person. The implant can be coated with materials such as vinyl polymers, eg acrylic polymers, polyethylene and carbon fiber filled Teflon. Journal Bone and Joint Surgery, Volume 53A, No. 1, Page 101 (1971)
, G. Galante et al. describe sintered fiber metal composites as a basis for attaching implants to bone, and U.S. Pat. No. 3,808,606, issued May 7, 1974, Stainless steel and cobalt-chromium-molybdenum alloy prostheses with porous surfaces for ingrowth fixation are described. Also of general interest are: November 23, 1976
No. 3,992,725, “Implantable Materials and Devices and Methods for Stabilizing Body Implants,” issued Oct. 1975.
U.S. Patent No. 3,909,852, “Implantable Replacement Structure for At least a Portion of a Middle Ear Skeletal Chain,” issued on May 7th;
and U.S. Pat. No. 3,971,670, issued July 27, 1976, entitled "Implantable Structure and Method for Making the Same." In addition to these patents, various articles on bone ingrowth into porous materials can also be found in the literature. Typical papers include, among others, S.F. Fulbert's ``Addition of prostheses to the musculoskeletal system by tissue ingrowth and mechanical fixation,'' Journal of Biomedical Materials Research Symposium, Vol. 4. , p. 1 (1973); M. Spector et al., "Bone Growth into Porous High Density Polyethylene", Journal of Biomedical Materials Research, Symposium, Vol. 7, No. 595
(1976); C.A. Homsey, "Implant Stabilization - Chemical and Biomedical Considerations", Orthopedic Clinics of North America,
Volume 4, No. 2, Page 295 (1973);
N. Kent et al., “Proplast in tooth surface reconstruction”, Oral Surgery, Oral Medicine, Oral Pathology, Vol. 39, No. 3, No.
347 pages (1975). However, porous materials disclosed in the literature as useful for prosthetic devices provide an unsuitable biomechanical environment leading to either of two undesirable conditions. First, low modulus-high creep porous coatings such as porous Teflon/graphite composites, after long periods of time,
It shows a metastable fibrous structure within the pores. This tissue is not suitable for supporting a force-bearing bonded prosthesis. Fibrous tissue is a metastable precursor to the skeleton and is remodeled into bone under normal physiological conditions (including physiological loading conditions). High loads transmitted through low modulus materials and excessive creep result in fibrous tissue that does not remodel into bone. Other low modulus-high creep materials used in prosthetic devices include polyethylene and polypropylene. Second, high modulus materials such as ceramics (16×10 ⁶ psi) and metals like titanium (17×
10 ⁶ psi) and cobalt-chromium-molybdenum alloy (34×10 ⁶ psi) do not exert sufficient stress on the ingrowth or surrounding skeleton to prevent reabsorption. In porous metal and ceramic coated femoral and humeral shafts, loads are concentrated at the apex of these prosthetic components, causing stress concentrations in the surrounding skeleton and subsequent resorption. Furthermore, the bone fragments within the pores of these porous ceramic and metal implants are not loaded and therefore do not resorb. Bone loss from the pores in the unloaded porous implant area is demonstrated histologically. This type of bone loss results in a reduction in composite strength (eg, interfacial shear strength), which in turn results in a reduction in "in-use" performance in these high modulus porous materials. The patents and publications cited above describe the use of porous coatings for prostheses and further describe requirements for acceptable pore size ranges. However, metals, ceramics, and polymers that have been disclosed as useful for coating prosthetic devices, such as vinyl polymers, polyethylene, polypropylene, and carbon-filled Teflon, provide adequate early fixation, long-term stability, and bone stability. It has been found that an appropriate biomechanical environment is not established to achieve strength at the prosthetic interface. In addition, previously described polymeric materials have sufficient toughness, creep resistance, and
May lack tensile and impact strength and steam sterilization. Even selected high density polyethylene and polypropylene porous compositions that are described as having reasonable flexibility and strength do not meet the general acceptance criteria. More recently, researchers have been able to achieve long-term bone fixation by having a coating of porous bionic thermoplastic materials, such as polysulfone, and by tissue ingrowth and subsequent remodeling of the bone after implantation. A prosthetic device was devised. This kind of device is
Described in US Pat. No. 4,164,794. There are
It has been shown that the coating can be applied either by sintering techniques or by foaming techniques to give the coating certain physical properties that optimize tissue ingrowth and bone remodeling and long-term bone fixation. has been done. It has now been found that improved results can be obtained in prosthetic devices having sintered coatings of biomedical thermoplastic materials by the presence of a porosity gradient across the coating thickness. The higher porosity on the outside facilitates bone ingrowth, while the lower porosity on the inside provides better attachment to load-bearing components and greater stiffness and strength. Accordingly, one or more of the following objectives may be achieved by implementing the present invention. It is an object of the present invention to provide a prosthetic device consisting of an inner load-bearing functional component and an outer sintered coating of a selected bionic thermoplastic having a porosity gradient over at least a portion thereof. It is to be. Another object of the present invention is to provide a coated prosthesis that achieves long-term bone fixation by tissue ingrowth into the selective porous bionic thermoplastic coating after implantation and subsequent remodeling into the bone. The purpose is to provide equipment. Furthermore, the object of the present invention is to
It is an object of the present invention to provide a prosthetic device with a coating having a specific porosity and porosity gradient that provides an optimal substrate for tissue ingrowth. Another object is to provide a prosthetic device in which the coating exhibits sufficient tensile and impact strength during and after bone formation to absorb applied loads during insertion and after surgery. Another object of the invention is to provide a porous bionic thermoplastic coating containing additives to improve biological and/or mechanical properties. Yet another object of the present invention is to provide a porous bionic thermoplastic coating containing additives to increase abrasion resistance. Another object is to provide one or more methods for manufacturing coated prosthetic devices or anatomically shaped structures made of bionic thermoplastics with porosity gradients. These and other objects will be readily apparent to those skilled in the art from the following description. Broadly speaking, the present invention relates to a porous bionic thermoplastic having a porosity gradient that allows tissue ingrowth into the covering material to firmly and permanently secure the device to the musculoskeletal system. Directed to a prosthetic device covered by the material. In one embodiment, the prosthetic device has a load-bearing
The thickness of the functional component and at least a portion of it is approximately
a porous coating of 0.5 to about 10 mm of bionic thermoplastic material, said material accommodating and guiding the ingrowth of reticular and dermal bone fragments, said coating having the following properties: (a) a substantial portion of the coating has an average pore size of about 90 to about 600 microns; (b) the pore connections have an average diameter of greater than about 50 microns; About 250,000 to about 250,000 to about a solid non-porous thermoplastic material
500000psi and about 500000 to about 500000 for reinforced solid non-porous thermoplastic materials
3,000,000 psi; (d) the total porosity is greater than about 20%, where there is a porosity gradient across the coating, with the smallest porosity on the side of the coating that contacts the load-bearing functional component; (e) the total creep strain of the unreinforced solid nonporous thermoplastic material is less than 1% at ambient temperature under a constant stress of 1000 psi; Something. All of these properties can transfer stress on the musculoskeletal system to bone fragments within the pores of the material, and
Sufficient to maintain sufficient loading and pore stability to promote irreversible ossification. Prosthetic devices manufactured by the method of the present invention exhibit well-defined directional porosity gradients, providing improved biomechanical performance over that of US Pat. No. 4,164,794. In particular, a preferred higher level of porosity and a larger average pore size are present on the outer surface of the covering that has to give rise to bone ingrowth; increased material density (lower porosity) i.e. Greater stiffness and strength exist near the inner surface of the coating where stress levels are highest; a smoother inner surface provides greater contact area with the metal implant and improves coating/metal bond strength. A surface is obtained. Optimally reduced processing times can be achieved. Because the process to obtain a well-defined porosity gradient relies on relatively high directional heating rates, this condition must be avoided in methods used to produce uniform porous parts. It is from. As shown in US Pat. No. 4,164,794, the materials used in the manufacture of prosthetic devices are classified as "bionic thermoplastics." One important feature of these materials is that
Their performance can be predicted over both long and short term periods through the use of metal design engineering equations. These engineering design equations apply only up to the linear viscoelastic limit of the material. High density polyethylene has a linear viscoelastic limit of less than 0.1%, and this limit on the magnitude of strain minimizes the stress that can be tolerated. In contrast, within the definition of this disclosure, the linear viscoelastic limit for bionic thermoplastics is at least 1% strain. For example, one preferred bionic thermoplastic material found suitable for the coating of the present invention is polysulfone, which has a 2% strain limit.
Therefore, both long- and short-term metallurgical design equations can be applied up to this limit. The unique properties of bioengineering thermoplastic materials become even clearer when their performance is compared to polymeric materials previously disclosed to be useful in porous fixation devices. If the creep modulus changes significantly over time, the deflection increases significantly, resulting in hole distortion and micro-displacement of the prosthesis under load. Creep tests have already been reported in the literature for porous high density polyethylene and polytetrafluoroethylene-graphite composites, both of which are shown in the patent specifications cited above. Regarding the porous polytetrafluoroethylene-graphite composite
Significant changes in pore structure were observed to occur at compressive stresses as low as 80 psi, and for porous high density polyethylene at 300 psi. Typical failure times for two reported high density polyethylene fabrications were less than 5 minutes when subjected to stress levels greater than 300 psi. It should be noted that this is a stress level not experienced in orthopedic bond and device applications.
The importance of maintaining pore geometry under loading environments has long been demonstrated in areas where fibrous tissue has been observed to develop within the pores. This is a surgical procedure that precedes bone ingrowth when the porous polymeric coating on the bonded prosthesis must have sufficient strength and stiffness to support the load independently without the assistance of the ingrowth bone. This is especially important in the early stages. The strength of conventional polymeric materials derives from the ingrowth of long bone. The bionic thermoplastic porous coating has a strength similar to bone. Each of these materials produced in accordance with the teachings of the present invention provides a coating with the physical properties listed above. Examples of these materials are polysulfones, such as polyphenylsulfone, polyethersulfone, polyarylsulfone; polyphenylene sulfide, polyacetals, thermoplastic polyesters, such as aromatic polyesters, polycarbonates, aromatic polyamides;
These include aromatic polyamideimides, thermoplastic polyimides, polyaryletherketones, polyarylethernitriles, and aromatic polyhydroxyethers. The most preferred materials for use in this invention are aromatic polysulfones. These polysulfones contain repeating units with the formula [Ar-SO ₂ ], where Ar is the structural formula a divalent aromatic compound containing at least one unit having the formula: wherein Y is oxygen, sulfur or the residue of an aromatic diol, such as e.g. 4,4'-bis(p-hydroxyphenyl)alkane. It is a family group. Particularly suitable polyarylene polyether polysulfone thermoplastics have the structural formula wherein n is equal to from 10 to about 500. These are UDEL polysulfone P-1700 and P-3703
It is commercially available from Union Carbide Company. These materials differ in that P-3703 has a lower molecular weight. Also useful are Astrel 360, a polysulfone sold by 3M, and Polysulfone 200P, sold by ICI, and Radel polyphenylsulfone, sold by Union Carbide. Certain crystalline bionic thermoplastics are also useful, such as styrane from Raychem, polyarylene and phenoxy A from Union Carbide. The materials used in the present invention can also contain reinforcing agents, if desired. Incorporation of glass, carbon or organic fibers into bionic thermoplastics improves their load-bearing and structural properties. 250,000 for bioengineering thermoplastics
It exhibits bulk tensile or flexural modulus values in the range of ~500,000 psi. Fiber-reinforced products exhibit modulus values of up to 3 million, depending on the fiber type and loading. These values of elastic modulus provide an intermediate range required for initial post-operative support and long-term stability of the implanted prosthesis in high load areas fixed by bony ingrowth. A bionic thermoplastic coating produced by the method of the invention and having a porosity gradient across the thickness of the coating will be apparent from the accompanying drawings. Referring to FIG. 1, the interior surface of a porous bionic thermoplastic coating made in accordance with the teachings of the present invention is shown at 75x magnification. At the interface 10 between the porous coating 12 and the solid support 14, the density of the sintered parts is greatest. These sintered surface particles form a substantially skin or continuous coating, which provides maximum surface area in contact with the support. Improved adhesion to the substrate is thus obtained while providing stiffness and strength where needed. Particles 16 on the side far from the interface 10
has a sintered shape characterized by larger pores and interconnections, i.e. maximum coverage porosity. FIG. 2 is also a photograph showing a cross-section of the porous polysulfone coating at 75x magnification, showing that the outer porosity of the coating is greatest at the outer surface 18, thus facilitating bone ingrowth. In practice, the prosthetic device of the invention is conveniently manufactured by sintering techniques, where the particles of bionic thermoplastic material undergo sintering, i.e. particle fusion at one or more points of contact. , heated in a specific sequence for a time and at a temperature sufficient to provide a porous continuous composite material of bioengineering thermoplastic having the desired porosity gradient and mechanical properties. It has been observed that, with respect to desired mechanical properties, the modulus of porous materials can be predicted by the Kellner equation or by the modified Halpin-Zwei equation. Thus, for example, to obtain a material with a porosity of 55% and a modulus greater than 40,000 psi, the modulus of the starting polymer must exceed 200,000 psi.
For example, most polypropylene and all high density polyethylenes cannot be processed into materials with porosity of 55% and modulus of 40,000 psi. On the other hand, the modulus of solid polysulfone is
In excess of 340,000 psi, a material with a porosity of 55% and a modulus of greater than 70,000 psi can be obtained. For prosthetic devices of the present invention, preferably the coating has a porosity of about 20% or more, more preferably from about 30 to about 70%.
% porosity. Although it has been possible to predict the modulus of thermoplastics with desired porosity, there has been no simple method by which materials can be processed that come close to these predicted values to be useful in the apparatus of the present invention. However, it has been unexpectedly discovered that by appropriate selection of particle size, molecular weight distribution, and sintering conditions, the desired degree of porosity can be obtained without sacrificing mechanical properties. All three conditions mentioned above are interrelated and necessary to obtain a coating or article with the required properties. For example, sintering times and temperatures that produce the desired pore size distribution may not produce the desired modulus and/or tensile strength. Starting particle size distribution and sintering time and temperature must be adjusted to achieve the desired balance of pore size, porosity, and mechanical properties. Regarding particle size distribution, sintered materials that meet the porosity and mechanical property requirements necessary for a successful prosthetic device can be manufactured using either a single particle size or a binomial distribution of particle sizes. can do. In practice, the particle diameter ratio is between about 7:1 and about 5:1.
A mixture of particle sizes such that the particle size is in the range of thickness 2
It has been found to be acceptable for porous coatings larger than mm. Particularly preferred are particle sizes of about 300μ to about 50μ. For example, 50 mesh sieve (US standard sieve)
A mixture of particles retained above and passed through a 270 mesh sieve provided coatings and articles with the desired porosity and biomechanical characteristics. Additionally, it has been observed that optimal results are obtained when the ratio of fine to coarse particle size ranges from about 40 to about 60% by weight. For porous sintered coatings on the order of 0.5 to about 2.0 mm thick, particles that pass through a 40 mesh sieve and are retained on a 70 mesh sieve (average particle size 320μ) have the most advantageous mechanical and biological It was found that it gives properties. As mentioned above, sintering conditions are particularly important in achieving not only the desired overall porosity of the coating, but also the desired porosity gradient in the same operation.
Generally, sintering involves loading powder into a metal mold, and then moving the mold to a predetermined sintering temperature Ts above the glass transition temperature Tg and below the melting temperature or melt processing temperature Tm (i.e., Tg<Ts<Tm). This was achieved by heating up to. This sintering temperature is kept constant for a predetermined time t. Virtually no pressure is applied other than that created by differential thermal expansion. Application of pressure at Ts results in melting of the material. This indicates that in the presence of pressure, lower temperatures and shorter time cycles must be used to maintain porosity in the sintered part. However, in contrast to the sintering technique disclosed in U.S. Pat. A high heating rate with directional heat flow is used, thus providing a well-defined porosity gradient. By using high heating rates, a temperature gradient exists within the sintered powder resulting in lower porosity in areas of higher temperature. Thus, by using directional heating, ie by passing the heat from the inside to the outside during the sinter bonding process, a regular porous structure is obtained. Depending on the particular prosthetic device being coated, one or more techniques can be used to provide differential heating. For example, a sleeve of sintered porous material can be manufactured separately from the prosthesis and secured to the device by gluing and/or shrinkage. As shown in Example 2 below, sleeves of sintered material are conveniently manufactured in insulated molds, where the core is heated to first form a "skin" on its inner surface, which skin It closely conforms to the shape of the mold and has low porosity adjacent the interface. Heating can be accomplished in several ways, such as by passing heated silicone oil through the interior of the core. After heating the inside is completed, the insulation is removed and the mold is heated both inside and outside or just the outside. This continues until sintering is complete and the material has the desired pore structure and mechanical properties. Once the material has cooled sufficiently, it is removed from the core mold. A separate sintered sleeve is then dip coated onto the implant with a thin film of PSF-silyl reactant (PSF-SR);
It can then be applied to metal implants by crosslinking at 270°C for 10-15 minutes.
The implants were then similarly coated by dip coating.
Apply a thin film of PSF-1700M.G. and dry it. Once this thin film dries, the cadre is reattached to the PSF−
1700M.G. solution (5-10% PSF− _{in CH2CH2}
1700 M.G.) and slip the porous sleeve into place while the PSF "glue" is still viscous. The finished prosthesis is then dried and cleaned. After sterilization, it is ready for surgical implantation with an interference fit. Other methods can also be used, especially if the implant is non-uniform or too small to be heated from the inside. For example, as shown in Example 3 below,
It has been observed that some human hip bone prostheses have dimensional variations between the same models. Induction heating can be used instead of machining a new model for each implant. The main differences between this technology and the prior art are:
The implant itself is used as the "male" or core part of the mold, and the powder is heated first from the outside and then from the inside. Sintering from the inside is achieved by induction heating the implant. The implant trunk is prepared for bonding by applying thereto a thin film of PSF-SR and curing, followed by applying and drying a thin film of PSF-1700M.G. The implant is then inserted into the mold cavity and secured. The gap (1 mm or more in width) between the implant trunk and the mold cavity wall is then filled with PSF of the desired particle size distribution.
Fill with 1700M.G. powder. The outside of the mold is then heated to the sintering temperature, for example by band heaters, electrical resistance heater tape, or induction coils if the mold is made of a suitable material. External heating is maintained only for a short time. This results in an oversized preform that is undersintered. This is because the polymer powder sinters maximally at the cavity surface and shrinks away from the implant. After the preform has sufficient integrity to maintain its shape, the cavity portions of the mold are separated and ejected.
This leaves an oversized covering on the implant.
The implant and coating are then placed within an induction coil designed to provide uniform heating of the metal implant surface. The implant is induction heated to the appropriate temperature for sintering and maintained at this temperature until sintering is complete and the coating contracts to conform to and bond to the surface of the implant. The coated prosthesis is then cooled and cleaned. After sterilization, it is ready for surgical implantation. The porous sintered coatings of the present invention having a porosity gradient are also useful in the root core of dental implants. However, the small size of the implant made it impractical to heat it only from the inside. Therefore, the mold was designed to provide a sufficiently large temperature gradient. The mold consisted of a top plate, a mold body, and a bottom plate. The base plate has one or more cavities for accommodating the implant to be coated in an inverted position. That is, the implant is
It is installed with the contact area to be coated facing upward. The mold body is then placed on top of the bottom plate, which has openings aligned with the upright posts. Thereafter, the gap between the core or post of the mold body and the cavity wall is filled with a polymer powder having the desired particle size distribution. The powder is introduced through holes in the top of the mold body and tamped to ensure that all gaps are fully filled. The top plate is then placed in place and secured with cap screws. The mold wrapped with insulation material is placed in a press whose surface plate is preheated. 40°C from the desired sintering temperature on the top surface plate.
℃ and the bottom platen to 40 ℃ above the desired sintering temperature. These settings cause the titanium core to heat first and then maintain a higher temperature than the cavity wall of the mold body. As a result of this, the inner surface of the porous coating corresponds very closely to the outer surface of the core to which it has to be bonded.
After the mold reaches the desired temperature and remains there for the desired time, the mold is removed from the press and allowed to cool. The finished implant is removed from the mold and cleaned. After sterilization, the artificial tooth root is ready for surgical implantation. The load-bearing functional components of the prosthetic devices of the present invention can be constructed from a variety of metals and alloys known in the art. Although titanium and tantalum are the only pure metals considered safe for most internal applications, various alloys have also been found to be generally acceptable. Stainless steel, cobalt-based alloys, and titanium-based alloys are all tolerated by the human body, resistant to corrosion, and processed into desired shapes. For some applications, it is desirable to incorporate additives that increase the wear resistance of the prosthesis. Carbon fibers, graphite fibers, Teflon, and molybdenum disulfide are useful additives that provide wear-resistant bionics that are equal to or better than self-lubricating materials typically used in commercially available bonded prostheses. Provides thermoplastics for use. injection molding or e.g. acetabulum,
Compositions containing carbon fibers are suitable for machining glenoid prostheses, such as the tibia and glenoid components of total knee and shoulder replacements. The sintered coating of the present invention can be bonded to the load-bearing functional component by several methods. For example, silyl-reactive polymers such as silyl-reactive polysulfones are used to bond porous polymeric coatings to metal substrates. Silyl-reactive polysulfone (PSF-
SR) resin has three important characteristics. Primarily,
The presence of hydrolyzable silane end groups provides an inherent bonding ability to metal surfaces. Second, PSF−
SR resin has low melt (or solution) viscosity,
It significantly facilitates "wetting" during the formation of adhesive bonds. Third, these are polymeric adhesives that are not soluble in physiological fluids and therefore have no biological/toxicological effects when implanted. Other techniques can also be used, such as heat shrinking a partially sintered sleeve. Example 1 Effect of Sintering Conditions on Pore Size This experiment shows the effect of sintering conditions (i.e. particle size, time and bath temperature) without any attempt to obtain a porosity gradient. For this experiment, a simple mold was fabricated from a steel tube with an outer diameter of 3/8 inch. The tube was cut to 6 inches long and fitted with a threaded end plug. The wall thickness of the tube was approximately 0.038 inch. The resulting sintered plastic part has a diameter of
It was 0.300 inches and 6 inches long. this is,
This was found to be a convenient sample size for tensile characterization. PSF- with the particle size distribution shown in the table below
3703 powder was used. This material was sintered according to the following schedule: the powder was filled into a mold; the mold was immersed in an oil bath at 220° C. for various times ranging from 10 minutes to 30 minutes. Then the resulting diameter
The 0.300 inch rod was cut into 2.5 inch sample lengths. The interconnected pore size distribution was then determined by mercury intrusion porosimetry and the data are reported in the table. Characteristic pore size is expressed as the percentage of pores greater than or equal to 132μ. As the time at temperature was increased from 10 to 30 minutes, the number of pores greater than or equal to 132μ in diameter increased. However, the material was
If kept for longer than minutes, the resulting sample will no longer be porous. On the other hand, if the material is exposed to that temperature for less than 10 minutes, little or no sintering will occur. Thus, there is an optimum time at a given temperature and optimum temperature for a given particle size and molecular weight distribution to achieve the desired pore size.

【table】

Table Example 2 Effect of heating order on porosity In this experiment, male-female molds were machined from steel tubes of different internal diameters. The male part had an outside diameter of 0.25 inches and was attached inside a larger diameter female part that had an inside diameter of 0.50 inches. The mold is
The arrangement was such that heated oil could flow between the internal hollow portion of the male part and the entire mold which was heated from the outside by other means. The gap between the core and cavity portion of the mold was then filled with a polymer powder of the desired particle size distribution. The powder was not compacted as in standard powder-metallurgical sintering techniques. The outside of the mold was insulated and silicone oil at a temperature of 250°C was flowed into the male part of the mold for 15 minutes. Then remove the insulation and use electric heating tape to reach an outside temperature of 240℃
The outside was heated until it reached . These temperatures are approximately
It was maintained for 10 minutes. After cooling, the sintered sleeve was removed from the mold and was found to have a porosity gradient similar to that shown in FIGS. 1 and 2. Example 3 Sintering of a conformal porous coating by heating the implant itself Dimensional changes (0.050) found to exist between different human hip implants of the same model
~0.060 inch), it would be necessary to machine a new mold for each hip bone to be covered. This example shows a method for induction heating of the implant itself, which provides simultaneous sintering and bonding of the coating. In this way, it was possible to cover a prosthesis which, due to its geometry, would not be able to be applied integrally with a conformal pre-sintered sleeve. The main difference between this technique and the prior art is that the implant itself is used as the male or mold core part, and the powder is heated first from the outside and then from the inside. Sintering from the inside is accomplished by induction heating the implant. The implant trunk was prepared for bonding by applying a thin film of PSF-SR to it and curing it, followed by applying a thin film of PSF-1700M.G. and drying it. The implant was then inserted into the mold cavity and secured. PSF-1700M.G with the desired particle size distribution is placed in the gap (width 1 mm or more) between the implant trunk (core) and the mold cavity wall.
Filled with powder. The outside of the mold was then brought to the sintering temperature and maintained for a short period of time. This resulted in an oversized preform that was undersintered. This is because the polymer powder sintered best at the surface of the cavity and shrank away from the implant. Once the preform had sufficient integrity to maintain its shape, the cavity portion of the mold was separated and removed. this is,
An oversized covering was left on the implant. The implant and coating were then placed inside an induction coil designed to provide uniform heating of the metal implant surface.
The implant was induction heated to a temperature suitable for sintering and maintained until sintering was complete and the coating contracted to conform to and bond to the surface of the implant. The coated prosthesis was then cooled and cleaned. After sterilization, it can be surgically implanted. Place the graft in a hot air oven with the trunk facing up.
The sintering/shrinking/bonding process was evaluated by heating to 255°C. The preformed porous sleeve was then slipped into the thermal trunk and allowed to complete its sintering and shrink to a snug fit. Even without adhesive bonding, this porous coating was difficult to remove. Although the invention has been illustrated by the foregoing examples, it should not be construed as limited to the materials used herein;
Rather, the present invention relates to the inclusive scope as disclosed above. Various modifications and embodiments can be made without departing from the spirit and scope of the invention.

[Brief explanation of drawings]

FIG. 1 is a micrograph of a cross section of the porous polysulfone coating showing the inner surface, and FIG. 2 is a micrograph of a cross section of the porous coating showing the outer surface. 10... Interface, 12... Porous coating, 14...
Solid support, 16... particles, 18... outer surface.

Claims (1)

Claims: 1. Consisting of a load-bearing functional component and a porous coating over at least a portion thereof of a bionic thermoplastic material suitable for and inducing ingrowth of bone fragments, said material comprising polysulfone. , polyphenylene sulfide, polyacetal, thermoplastic polyester, polycarbonate, aromatic polyamide, aromatic polyamideimide, thermoplastic polyimide, polyaryletherketone, polyarylethernitrile, and aromatic polyhydroxyether and the following properties: (a) a substantial portion of the coating has an average pore diameter of from about 90 to about 600 microns; (b) the pore interconnections have an average diameter of greater than about 50 microns; (c) a modulus of elasticity of 250,000 to approx. for reinforced solid non-porous thermoplastic materials
500000psi and about 500000~ for reinforced solid non-porous thermoplastic materials
(d) has a total porosity of greater than about 20% and there is a porosity gradient across the coating, with the smallest pores on the side of the coating that contacts the load-bearing functional component; (e) The total creep strain of the unreinforced solid nonporous thermoplastic material is less than 1% at ambient temperature at a constant stress of 1000 psi; All of these properties maintain sufficient load and pore stability to transfer musculoskeletal stresses to the bone fragments within the pores of the material and promote irreversible ossification. A prosthetic device characterized in that it is sufficient to. 2. The device of claim 1, wherein the porous coating has a thickness of about 0.5 to about 10 mm. 3. The device of claim 1, wherein the thermoplastic material is reinforced and has a modulus of elasticity of about 500,000 to about 3,000,000 psi. 4. The device of claim 1, wherein the thermoplastic material is unreinforced and has a modulus of elasticity of about 250,000 to about 500,000 psi. 5. The device of claim 1, wherein the thermoplastic material has a porosity of about 30 to about 70%. 6. The device according to claim 1, which is a total hip bone prosthesis. 7. The device of claim 1 which is an endosteal blade implant or other dental implant. 8. The device according to claim 1, which is an intramedullary nail. 9. The device according to claim 1, which is a reticulated screw. 10. The device according to claim 1, which is a cortical screw. 11. The device according to claim 1, wherein the load-bearing functional component is made of stainless steel. 12. The device according to claim 1, wherein the load-bearing functional component comprises a cobalt-based alloy. 13. The device according to claim 1, wherein the stress-proof functional component comprises a titanium-based alloy. 14. The device of claim 1, wherein the bionic thermoplastic material is polysulfone. 15. The device of claim 1, wherein the bionic thermoplastic material is polyarylsulfone. 16. The device of claim 1, wherein the bionic thermoplastic material is polyphenylsulfone. 17. The device of claim 1, wherein the bionic thermoplastic material is polyethersulfone. 18. The device of claim 1, wherein the bionic thermoplastic material is polycarbonate. 19. The device of claim 1, wherein the bioengineering thermoplastic material is an aromatic polyamide. 20. The device of claim 1, wherein the bioengineering thermoplastic material is an aromatic polyhydroxyether. 21 consisting of a load-bearing functional component and a porous coating over at least a portion thereof of an engineered thermoplastic material that is biologically suitable and induces bone fragment ingrowth, said material including polysulfone, polyphenylene, etc. a prosthetic device selected from the group consisting of fluoride, polyacetal, thermoplastic polyester, polycarbonate, aromatic polyamide, aromatic polyamideimide, thermoplastic polyimide, polyaryletherketone, polyarylethernitrile, and aromatic polyhydroxyether. (a) a male-female mold in which the male part of the mold is in the shape of at least a portion of the load-bearing functional component, the gap between the core and the cavity portion comprising: ~
filling a particulate biomedical thermoplastic material having at least a portion of the average particle size distribution within the range of about 600 microns; (c) thereafter increasing the temperature of the female part of said mold to approximately the same temperature; (d) raising the temperature of at least one part to at least partially (e) further heating said material to complete sintering and providing said material with the following properties: (a) a substantial portion of the coating is between about 90 and about 600 microns (b) the pore interconnections have an average diameter of greater than about 50μ; (c) a modulus of elasticity of from about 250,000 to about 250,000 for unreinforced solid non-porous thermoplastic materials;
500000psi and for reinforced solid non-porous thermoplastic materials approx.
500,000 to about 3,000,000 psi; (d) has a total porosity of greater than about 20% and there is a porosity gradient across the coating, with the smallest pores on the side of the coating that contacts the load-bearing functional component; (e) the total creep strain of the unreinforced solid nonporous thermoplastic material is 1% at ambient temperature at a constant stress of 1000 psi; (f) fixing said material to said load-bearing functional component. 22. The method of claim 21, wherein the particulate bioengineering thermoplastic material has at least a portion of a binomial distribution of mean particle size ratios ranging from about 7:1 to about 5:1. 23. The method of claim 21, wherein the average pore size is from about 150 to about 400 microns. 24. The independent partially sintered material is removed from the mold after completion of step (d) and further heated to complete the sintering of said material.
The method described in section. 25. The method according to claim 24, wherein the sintered material is fixed to the load-bearing functional component by mechanical and/or chemical means. 26 consisting of a load-bearing functional component and a porous coating over at least a portion thereof of an engineered thermoplastic material biologically suitable and inducing ingrowth of bone fragments, said material including polysulfone, polyphenylene, etc. a prosthetic device selected from the group consisting of fluoride, polyacetal, thermoplastic polyester, polycarbonate, aromatic polyamide, aromatic polyamideimide, thermoplastic polyimide, polyaryletherketone, polyarylethernitrile, and aromatic polyhydroxyether. (a) a gap between the core and the cavity in a male-female mold in which the male part of the mold is an uncoated load-bearing functional component, the gap between about 50 and about 600 microns; filling a particulate bioengineering thermoplastic material having at least a portion of an average particle size distribution within a range; (b) adjusting the temperature of the female portion of the mold to be equal to the glass transition temperature of the thermoplastic material; (c) raising the female part of the mold to a temperature between the processing temperature and a time sufficient to sinter the material into a preform that retains its shape upon removal of the female part of the mold; removing, leaving an oversized partially sintered preform on the core; (d) then heating the core to conformally shrink and complete sintering of the preform; A method for manufacturing a prosthetic device, characterized by: 27. The method according to claim 26, wherein the stress-proof functional component is heated by induction heating. 28 consisting of a load-bearing functional component and a porous coating over at least a portion thereof of an engineered thermoplastic material that is biologically suitable and induces bone fragment ingrowth, said material including polysulfone, polyphenylene, etc. a prosthetic device selected from the group consisting of fluoride, polyacetal, thermoplastic polyester, polycarbonate, aromatic polyamide, aromatic polyamideimide, thermoplastic polyimide, polyaryletherketone, polyarylethernitrile, and aromatic polyhydroxyether. A method of manufacturing, comprising: (a) a male-female mold in which the male part of the mold is in the shape of at least a portion of the load-bearing functional component; ~
(b) filling the mold with a particulate bioengineering thermoplastic material having at least a portion of the average particle size distribution within the range of about 600 microns; (c) then heating said first surface plate to a temperature 40° C. higher than the sintering temperature of said material; A method for manufacturing a prosthetic device, characterized in that the second surface plate is heated to a temperature that is approximately the same number of degrees lower than the sintering temperature.