JP3085903B2 - Implant material for living hard tissue and method for producing the same - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to an implant material for living hard tissue and a method for producing the same, and in particular, it has excellent biocompatibility, has fracture toughness that can withstand practical use, and has affinity for living tissue. The present invention relates to an implant material for living hard tissue suitable as a substitute material for living hard tissue such as an artificial bone and an artificial tooth root, and a method for producing the same.

[0002]

2. Description of the Related Art Surgical treatments for repairing bone defects by embedding artificial bones made of metal or ceramics into bones lost due to accidents or diseases have become widespread. In addition, the artificial tooth root is embedded as an implant material in the tooth part that has come off due to tooth decay, etc., as a base, and crystallized glass is laminated on this base to form a denture, or made of metal or ceramic Restoration methods for fixing dentures are also widely used.

As properties required for implant materials for living hard tissues such as artificial bones and artificial tooth roots, it is essential that, as with general implant materials, firstly, they have excellent compatibility with living tissues. In other words, it has no harm to the body such as toxicity, irritation and carcinogenicity, does not induce blood coagulation, hemolysis or abnormal metabolism, and does not cause dissolution, corrosion or deterioration even under severe conditions in the living body. It is required to be chemically stable.

[0004] It is also important as a second required characteristic to have a structural strength and a fracture toughness sufficient to withstand repeated fluctuating loads. Furthermore, it is also an important required characteristic to have a property of being sufficiently compatible with surrounding bone and soft tissues after implanting in a living body as a third required property, so-called tissue affinity (biocompatibility).

[0005] Bone of a living body contains calcium-deficient carbonate ion-containing hydroxyapatite fine particles as a main inorganic component, and these hydroxyapatite fine particles and collagen fibers form a three-dimensional composite structure to form strong bone tissue. .

Conventionally, various metallic materials, organic materials, and inorganic materials have been used as implant materials for living hard tissues such as artificial bones and artificial tooth roots, which can be substituted for living bones. For example, metal materials such as stainless steel, titanium, titanium alloy, and cobalt-chromium alloy are used from the viewpoint of excellent mechanical properties, moldability, machinability, and impact resistance, and relatively low harm to living organisms. Have been.

However, since implants made of a metal material have no affinity for bone tissue at all, it is difficult to semi-permanently fix the implant to a bone defect, and there is a problem that durability is low. In addition, implants made of metal materials are dissolved in the living body depending on the material, especially the eluted heavy metal ions accumulate in specific organs and cause dysfunction of the organs, or wear out during use to generate metal fine particles, There is also a problem that the surrounding soft tissue is adversely affected. Metal implants are commonly used with bolts,
Although it is fixed to the bone with a nut, if the bone is thin, a gap is likely to be formed, and there is also a technical problem to be solved in that the tightly joined state between the implant and the bone tissue cannot be maintained for a long time.

[0008] As an organic implant material,
Various types of simple plastics and organic plastic materials reinforced with organic fibers or inorganic fibers are used. However, the organic implant material has a disadvantage that it is easily oxidized in a living body or deteriorated by hydrolysis. In addition, there is a problem in that the monomer is easily eluted and adversely affects the living body, and further, there is a problem that there is no affinity for bone tissue.

On the other hand, among the inorganic materials, hydroxyapatite (HAP) has a chemical composition of Ca ₁₀ (PO ₄ ) ₆ (OH) ₂ and is a calcium-deficient carbonate ion which is a main component of living bone. It has basically the same crystal structure as the contained hydroxyapatite, and is excellent in biocompatibility and bone tissue affinity. However, HAP ceramic implants produced by pressure molding and sintering of HAP fine particles have insufficient mechanical properties, and particularly have a fracture toughness value K _IC of about 0.69-1.16 MPa m ^1/2. However, since it is lower than the fracture toughness value (2.2 to 4.6 MPam ^1/2 ) of the dense bone of a living body, there is a problem that the bone is easily broken by a slight impact force. For this reason, implants made of HAP alone have not been put to practical use except for special applications.

[0010] Among the inorganic materials, zirconia sintered bodies, particularly partially stabilized zirconia (PSZ), are less harmful to the living body and have a fracture toughness value of 5.7 to 9.6 MPa.
^Due to its high mechanical strength of about ^1/2, it has attracted attention as a suitable implant material. However, since the above-mentioned partially stabilized zirconia has no affinity for bone tissue, it cannot be expected at all that the partially stabilized zirconia firmly fuses with bone tissue after implanting.

Thus, conventional metal materials, inorganic materials,
An implant made of an organic material has all the required characteristics that it has mechanical strength equal to or higher than that of living hard tissue such as bone, is compatible with bone tissue and surrounding soft tissue, and has good affinity for bone tissue. No satisfactory material has been obtained.

On the other hand, the above hydroxyapatite (HA)
In order to combine the properties of P) and the properties of partially stabilized zirconia (PSZ), HAP fine particles and PSZ
Attempts have also been made to develop an implant material having excellent affinity for bone tissue and excellent mechanical properties by mixing fine particles with the obtained fine particles, and pressing and sintering the obtained mixture.

[0013]

However, the above H
In the case where the AP fine particles and the PSZ fine particles are uniformly mixed by a normal mixing method, and the obtained mixed powder is pressed and then sintered to form an implant material, the mixing of both components at the fine particle level is difficult. It was incomplete, and two components were easily separated in the mixed powder, and it was difficult to improve mechanical properties as expected.

That is, according to the conventional mixing method, in the state of a mixed powder, even if a state in which both components are uniformly mixed is obtained, re-separation of the two components occurs in a subsequent step. As shown in the cross section of the sintered body in FIG. 6 (lower figure), the HAP fine particles 30 easily aggregate or combine in a chain to form the continuous phase 50. In this case, since the brittle continuous phase 50 composed of hydroxyapatite having a relatively low strength as compared with the zirconia sintered body is formed, the strength varies, and the fracture toughness of the implant material as a whole does not improve. There is a problem that only an implant material having poor durability can be obtained.

Further, as one component of the mixed powder, the particle diameter is 1 μm.
When an implant material is obtained by sintering using the following fine HAP particles, there is no HAP particle having a sufficiently large area on the surface of the implant material. There was also a problem to do. That is, after implant,
Osteoblasts create new bone between the bone tissue and the implant material such as artificial bone to bind the two tissues together. However, since there is no HAP tissue having a sufficiently large area, a strong bond is formed. There was a drawback that it could not be realized.

Further, when manufacturing an implant material comprising a conventional HAP / PSZ binary sintered body, the HA
When sintering is performed at a high temperature in a state where the P component and the PSZ component coexist, a part of zirconia (PSZ) changes from a tetragonal system to a cubic system, and a part of the zirconia (PSZ) changes with the phase change. HAP changes to tricalcium phosphate (TCP). It has been clarified by the present inventors that this phenomenon is accelerated as the particle size of both components becomes smaller. This change leads to a decrease in mechanical strength and bone tissue affinity of the implant material. In any case, an implant material having both excellent mechanical strength and biological tissue affinity has not been obtained.

The present invention has been made to solve the above-mentioned problems, and has excellent biocompatibility, has fracture toughness that can withstand practical use, has affinity for living tissue, and has an artificial bone or artificial bone. An object of the present invention is to provide an implant material for living hard tissue suitable as a substitute material for living hard tissue such as a tooth root and a method for producing the same.

[0018]

Means for Solving the Problems In order to achieve the above object, the present inventors have developed various ceramic powders having no mechanical harm and excellent mechanical properties, particularly excellent fracture toughness, and biocompatible fine particles. Evenly mixed by various mixing methods,
Furthermore, implant materials having various structures are prepared by molding and sintering, and the mixing method, raw material particle shape,
The effects of particle size and sintered compact structure on the fracture toughness of implant materials were compared by experiments.

As a result, a ceramic powder such as zirconia or alumina having a small particle size and a spherical hydroxyapatite powder having a relatively large particle size are compounded by a high-speed impact method in an air stream. It has been found that an implant material having excellent fracture toughness can be obtained when pressure sintering is performed. From observation of the cross-sectional structure of the sintered body, it was strongly inferred that the area occupied by the HAP was large and the affinity for living tissue was excellent. Particularly when forming a continuous phase consisting of a ceramic sintered body having a high toughness value and a material structure in which hydroxyapatite particles having relatively low toughness are discontinuously dispersed in the continuous phase,
It has been found that an implant material having a high toughness value can be obtained as a whole material. The present invention has been completed based on the above findings.

That is, the implant material for living hard tissue according to the present invention is characterized in that it has a microstructure in which particles having biocompatibility are dispersed in a continuous phase composed of a high toughness material and present as a discontinuous phase. . Further, the high toughness material comprises at least one of partially stabilized zirconia and alumina. On the other hand, the biocompatible particles are characterized by comprising at least one selected from hydroxyapatite, tricalcium phosphate and bioglass.
Further, when the weight ratio of the high toughness material such as a ceramic sintered body is 20% or more, preferably 50% or more, a high toughness implant material can be obtained.

It is desirable that the average particle size of the high toughness material particles is 1/5 or less of the average particle size of the biocompatible particles. Further, it is preferable that the biocompatible particles be exposed on the surface of the implant material. This implant material for living hard tissue is suitable as an artificial bone or an artificial tooth root.

Further, according to the method for producing an implant material for a living hard tissue according to the present invention, there is provided a method for manufacturing a living hard tissue having a fine structure in which particles having biocompatibility are dispersed as a discontinuous phase in a continuous phase formed of a tough material. In the method for producing an implant material for tissue, a composite particle is prepared by immobilizing a high toughness material powder by a high-speed air current impact method so as to coat the surface of a biocompatible particle with the high toughness material powder. The method is characterized in that the particles are subjected to pressure molding, and the obtained molded body is sintered in a non-oxidizing atmosphere.

In the above manufacturing method, the average particle size of the high toughness material powder is preferably set to be not more than 1/5 of the average particle size of the biocompatible particles. Further, the pressing operation and the sintering operation of the composite particles may be simultaneously performed by a hot press method.

Examples of the high toughness material used to construct the above implant material include zirconia (ZrO ₂ ) and alumina (Al
₂ O ₃ ) is preferred, and in particular, partially stabilized zirconia (PSZ) having a high toughness value is preferred. Further, as the high toughness material, a metal material such as stainless steel, Co—Cr alloy, Ti, Ti alloy, and tantalum (Ta) may be used.

As the particles having biocompatibility, there can be used hydroxyapatite (HAP), tricalcium phosphate (TCP), bioglass, etc., which are easily fused with the living tissue. In particular, hydroxyapatite particles have the same crystal structure as living bone tissue, and are most desirable for imparting bone tissue affinity to implant materials. The surface of the hydroxyapatite particles is partially changed to α-tricalcium phosphate (α-TCP) depending on the conditions during the sintering operation. This production of α-TCP relatively reduces the amount of hydroxyapatite that gives bone tissue affinity,
The sintering conditions for suppressing the generation of α-TCP as much as possible to cause peeling at the interface between the AP and the generated α-TCP and reduce the bonding strength between the implant material and the surrounding tissue. Is preferably selected.

The average particle diameter of the biocompatible particles such as hydroxyapatite is preferably in the range of 1 to 100 μm. When fine particles having an average particle size of less than 1 μm are used, the solid phase reaction easily proceeds in the sintering process, the HAP particles are easily converted to α-TCP, and the hydroxyapatite has a large area on the surface of the implant material. , It is not possible to obtain conformability with the surrounding tissue (healing strength).

On the other hand, when the average particle diameter is increased so as to exceed 100 μm, particles having biocompatibility are likely to be pulverized by an impact force in a step of immobilizing a high toughness material powder by a high-speed air impact method described below. In any case, it becomes difficult to impart a predetermined bone tissue affinity to the implant material. Therefore, the average particle size of biocompatible particles such as hydroxyapatite is 1 to 100 μm.
m, but a range of 5 to 50 μm is more preferable.

On the other hand, the average particle size of the high toughness material powder such as zirconia is determined by uniformly embedding the high toughness material powder on the surface of the biocompatible particles in the immobilization treatment step by the high-speed airflow impact method described later. In order to compound
The average particle size of the particles having biocompatibility is set to 1/5 or less. When the average particle size of the high toughness material powder is 1/5 or less of the average particle size of the biocompatible particles, the high toughness material powder is fixed so as to uniformly cover the surface of the biocompatible particles. And it is possible to prepare good composite particles.

Particularly, spherical hydroxyapatite particles are:
Primary particles having a particle size of about 0.5 μm are formed as a group, and are formed as base particles having a large number of irregularities on the surface. By fixing the high toughness material powder on the surface of the base particles, the high toughness material powder as child particles is buried in the uneven portions of the apatite particles, and both are integrated with a high bonding strength by a so-called mechanical anchoring effect. . In this case, the average particle size of the high toughness particles is 1/5 or less of the hydroxyapatite particle size, more preferably 1/10.
By performing the following, the high-toughness particles as the child particles can be uniformly coated on the entire surface of the apatite particles as the base particles, and a composite particle having a high two-component bonding strength can be obtained.

The mixing ratio between the high toughness material and the material having biocompatibility in the implant material can be arbitrarily set, but the weight ratio of the material having biocompatibility is 10 to 80%.
Should be set in the range. Although the strength characteristics of the implant material are improved as the compounding ratio of the high toughness material is increased, the compounding ratio of the particles having biocompatibility is relatively reduced, so that the bonding strength with the surrounding biological tissue is reduced. become. Incidentally, the partially stabilized zirconia as the high toughness material and the hydroxyapatite as the biocompatible particles are 10:10 or more (the ratio of the high toughness material is 50%).
%), The fracture toughness value of the implant material of the present invention is 2.8 MPam ^1/2 or more, and this value is 2.5 times or more of the fracture toughness value of hydroxyapatite alone. Fracture toughness value of compact bone
The value is equal to or greater than 4.6). The weight ratio is 1
The fracture toughness value of the implant material in the case of 0: 2 was 1.8 MPam ^1/2 , which was 1.6 times or more that of hydroxyapatite alone.

The high toughness material powder and the biocompatible particles serving as the raw material of the implant material are heated (calcined) at a temperature of 800 to 1100 ° C. in advance for 1 to 3 hours as a pretreatment, so that the surface thereof is formed. It is preferable to remove the adsorbed moisture and volatile impurities, and to maintain a structural strength that is not destroyed even by an impact force at the time of immobilization treatment by a high-speed airflow impact method described later.

Next, the mixed powder of the high-toughness material powder and the biocompatible particles heat-treated as described above is compounded by a high-speed air-flow impact method, and the surface of the biocompatible particles is coated with the high toughness. The process proceeds to a step of preparing composite particles by performing an immobilization treatment so as to cover the material powder.

FIGS. 1 and 2 are a system diagram and a cross-sectional side view of a main part showing an example of the structure of a high-speed airflow impact type powder surface reforming apparatus (hybridizer) 1 used for producing the composite particles. FIG. This powder surface reforming device 1
While repeatedly dispersing the raw material particles in the gas phase, mechanical energy mainly including an impact force that does not destroy the particles themselves and, if necessary, thermal energy are repeatedly applied to the raw material particles for 1 to 5 minutes. It is a device developed to efficiently and efficiently perform the immobilization process and the film formation process of different materials in a short time.

The powder surface reforming apparatus 1 includes a main casing 2, a rear cover 3 and a front cover 4, a rotor 5 rotating at high speed in the casing 2, and a predetermined interval between the outer periphery of the rotor 5. And a hammer-type or blade-type impact pin 6 radially provided around
A rotating shaft 7 rotatably supporting the inside of the casing 2 and a collision ring 8 which is provided along the outermost raceway surface of the impact pin 6 and is provided around the impact pin 6 with a certain space. It is configured with.

As the collision ring 8, rings of various shapes such as a concavo-convex type or a circumferential flat type are appropriately used.
The gap between the outermost raceway surface of the impact pin 6 and the collision ring 8 is adjusted to 0.5 to 20 mm, depending on the type of the object to be modified and the processing capacity of the surface modification device. Further, an opening / closing valve 9 is provided which fits closely with a modified powder discharge port provided by partially cutting out the upper part of the collision ring 8. Further, a valve shaft 10 of the opening / closing valve 9 and this valve shaft 10 are connected to each other. An actuator 11 that drives and operates the on-off valve 9 through the intermediary is provided.

A circulation circuit 12 which connects an inlet 12a opening to a part of the inner wall of the collision ring 8 and an outlet 12b opening to the front cover 4 located at the center of the rotor 5 to form a closed circuit; A raw material hopper 13, a raw material supply chute 14 for communicating the raw material hopper 13 with the circulation circuit 12, and an on-off valve 15 provided in the middle of the chute.
And an impact chamber 16 provided between the outer periphery of the rotor 5 and the collision ring 8.

On the secondary side of the on-off valve 9, a modified powder discharge pipe 17 and a bag collector 18 are additionally provided. Also,
A known preprocessor 19 such as various mixers and an automatic mortar used when it is necessary to attach child particles to mother particles in advance, and the powder mixed by the preprocessor 19 is supplied to the surface reforming apparatus 1 in a constant amount. And a raw material measuring feeder 20 are provided.

The above-mentioned powder surface reforming apparatus 1 is designed so that a batch operation can be performed automatically by a time control device (not shown). When the surface reforming apparatus 1 is operated for a long time, or when the processing capacity is large, a powder separating device such as a cyclone may be provided instead of the bag collector 18, or a plurality of bag filters may be used. The modified powder is configured to be connected in parallel or to discharge the modified powder using an exhaust fan (not shown).

Since the powder surface reforming apparatus 1 shown in FIGS. 1 and 2 is a complete batch type reforming apparatus, the ambient temperature in the apparatus may increase as the powder processing time elapses.
Therefore, a jacket 21 is formed inside the collision ring 8,
It is preferable that the temperature inside the shock chamber 16 and the circulation circuit 12 can be controlled to be constant by flowing cooling water or the like through the jacket 21. Further, by previously replacing the interior of the impact chamber 16 and the circulation circuit 12 with an inert gas, it is possible to perform the immobilization of particles by the present apparatus in an inert gas atmosphere.

When the particles are immobilized using the powder surface reforming apparatus 1, the apparatus is operated in the following manner.

First, the on-off valve 15 disposed in the middle of the raw material supply chute 14 is closed, and the on-off valve 9 at the modified powder outlet is also closed. Drive, for example, at an outer peripheral speed of 60 to 100 m
The rotor 5 is rotated at about / sec. At this time, an abrupt air flow is generated with the rotation of the impact pin 6, and a fan effect based on the centrifugal force of the air flow causes the circulation circuit 1
An outlet 12 opening at the center of the front cover 4 via
A circulation flow returning from b to the shock chamber 16 is formed.

This air flow is a complete self-circulating flow, and the amount of circulating air generated per unit time at this time is extremely large compared to the total volume of the shock chamber and the circulating system. An enormous number of airflow circulation cycles are formed in time. For example, a rotor with an outer diameter of 118 mm
The circulation air volume when rotating at an outer peripheral speed of m / sec is 0.
48 m ³ / min, and the number of air circulation cycles per unit time is 774 times / min. Since the amount of circulating air is proportional to the outer peripheral speed of the rotor, the number of air circulation cycles per unit time also increases in proportion to the outer peripheral speed of the rotor.

Here, when the immobilization treatment performed by the method of the present invention is performed, the outer peripheral speed of the rotor is 30 to 150 m /
It is preferably set in the range of sec, more preferably in the range of 40 to 80 m / sec. When the outer peripheral speed is less than 30 m / sec, a sufficient airflow is not generated in the circulation circuit, and the processing time becomes longer particularly when particles having a large specific gravity difference are processed, and the sufficient amount of air required for fixing the particles It is difficult to apply a strong impact force, and the processing efficiency is reduced. On the other hand, 15
It is difficult to obtain an outer peripheral speed exceeding 0 m / sec because of limitations in the structure of the device and mechanical structure, and particles having biocompatibility as base particles are easily broken (crushed). The immobilization process is completed in about 1 to 5 minutes.

With the circulating flow formed as described above, the on-off valve 15 is then opened, and the mixed powder of the high toughness material powder and the biocompatible particles is supplied to the raw material hopper 13 via the measuring feeder 20. The mixed powder enters the shock chamber 16 via the raw material hopper 13 and the raw material supply chute 14 when the mixed powder is charged in a short time. After confirming that no mixed powder remains in the raw material hopper 13, the on-off valve 15
Is closed. When the automatic batch operation is performed, the time required for charging the mixed powder is measured in advance and input to the time control device.

The mixed powder introduced into the shock chamber 16 is instantaneously hit by an impact pin 6 attached to a rotor 5 rotating at a high speed in the shock chamber 16, and is further applied to a peripheral collision ring 8. collide. Then, it is entrained by the circulating airflow, returns to the shock chamber 16 again around the circulation circuit 12, and receives the same impact action again. By repeatedly receiving the striking action in this manner, a uniform immobilization process is performed in a short time of several tens seconds to several minutes. as a result,
The high-toughness material powder as child particles can be firmly immobilized on the surface of the particles having biocompatibility as the mother particles, and composite particles having the surface of the mother particles uniformly coated with the child particles are formed. In the case where the child particles are low freezing point substances, the child particles are melted at the moment of the impact action and are fixed on the surface of the base particles in a film form.

After the completion of the fixing process, the on-off valve 1
5 is opened, and the open / close valve 9 of the modified powder outlet is moved to the position shown by the broken line and opened to discharge the composite particles. The composite particles are subjected to the shock chamber 16 and the circulation circuit 12 in a short time of several seconds by centrifugal force acting on the particles themselves.
After that, the air is discharged from the apparatus main body, and further collected by the bag collector 18 via the discharge pipe 17.

Thus, the composite particles produced by the powder surface reforming apparatus 1 based on the high-speed impact in air stream have child particles simply adhered to the surface of the base particles like composite particles prepared by a conventional mixing apparatus ( It is not a so-called ordered mixture in a (glazed) state, but is a composite particle in which child particles are embedded in the surface of a base particle and firmly and densely fixed. Therefore, even when the obtained composite particles are subjected to pressure molding or heat treatment, the composite particles 2
It is less likely that the components are re-separated or that the components aggregate to form coarse particles.

Next, in the case where the high toughness material powder is left unfixed due to the relationship between the obtained composite particles and the compounding ratio, the mixed powder with the powder is pressed, then sintered and sintered to a predetermined shape. It moves on to the process of manufacturing an implant material.

As the above-mentioned pressure molding method, a uniaxial pressure (UAP) molding method using a usual mold pressing device or a cold isostatic pressure (CIP) molding for applying a molding pressure to a mixed powder isotropically. A law can be adopted. It should be noted that a hot press (HP) method or a hot isostatic pressure (HIP) method in which the above-described molding operation and sintering operation are performed simultaneously can also be used.

When sintering is performed after the above uniaxial pressure molding, first, the mixed powder is compacted at a molding pressure of 10 to 50 MPa for 2 to 5 minutes.
For 1 minute at a temperature of 1000 to 1450 ° C. in a non-oxidizing atmosphere such as argon gas.
Sinter for ~ 3 hours.

When sintering is performed after cold isostatic pressure (CIP) molding, first, a compacting pressure of 20 to 250 MPa is applied to the mixed powder for 3 to 6 minutes to form a compact by isostatic compression. , And the obtained molded body was similarly heated to a temperature of 1000 to 1450 ° C.
And sinter for 1-3 hours.

When the molding operation and the sintering operation are simultaneously performed by the hot press method, the raw material mixed powder is subjected to a pressure of 20 to 100 in a non-oxidizing atmosphere such as argon gas.
It is preferable to perform pressure sintering at a temperature of 1300 to 1400 ° C. for 5 to 20 minutes.

Among the various forming and sintering methods, according to the hot press (HP) method, a sintering temperature is relatively low, and a dense sintered body is easily obtained even if the sintering time is shortened. have. Therefore, while using hydroxyapatite (HAP) particles as biocompatible particles,
When sintering using partially stabilized zirconia (PSZ) as a high toughness material, a part of the HAP particles is suppressed from changing to tricalcium phosphate, and a part of the PSZ is tetragonal. From the cubic system to the cubic system.

As described above, the high toughness material powder and the particles having biocompatibility are compounded by a high-speed airflow impact method,
After the resulting composite particles are pressed and molded, the implant material prepared by sintering remains on the surface of the HAP particles with PSZ.
Because the particles are firmly fixed, the proportion of biocompatible particles exposed on the surface of the implant material is reduced, and if the material is used as it is as an artificial bone or artificial tooth root, there is less familiarity with the surrounding tissue Healing strength may decrease.

It is desirable that the surface of the obtained implant material is ground or polished to expose biocompatible particles to the surface, thereby increasing the biocompatibility of the implant material.

According to the implant material for living body hard tissue and the method for producing the same according to the above-described structure, strong composite particles obtained by compounding a high toughness material powder and particles having biocompatibility by a high-speed air impact method are used. Therefore, even when pressure molding or sintering is performed, it is less likely that the two components are separated again or that the components are aggregated to form coarse particles. Particularly, particles having biocompatibility are dispersed in a continuous phase composed of a tough material to form a microstructure existing as a discontinuous phase.

Therefore, it is possible to provide an implant material which is excellent in biocompatibility, has high fracture toughness and has affinity for living tissue, and is suitable as a substitute material for living hard tissue such as an artificial bone or an artificial tooth root. it can.

[0058]

DESCRIPTION OF THE PREFERRED EMBODIMENTS Embodiments of the present invention will be described more specifically with reference to the following examples.

Examples 1 to 3 and Comparative Example As particles having biocompatibility, the maximum particle size was 20 μm,
Spherical hydroxyapatite (HAP) particles (manufactured by Mitsubishi Materials KK) having an average particle size of 9.32 μm were prepared, and subjected to a heat treatment at a temperature of 800 ° C. for 2 hours as a pretreatment, whereby moisture adsorbed on the particles was obtained. In addition to removing volatile impurities, a calcined body was used to increase the structural strength.

FIG. 3 is a scanning electron micrograph showing the particle structure of the spherical HAP particles subjected to the above pretreatment. The spherical HAP particles are composed of an aggregate of fine primary particles having a particle size of about 0.5 μm, and it can be observed that a large number of irregularities are present on the surface. No change in the shape and size of the particles was observed before and after heating in the pretreatment. Also, HA
The chemical composition of the P particles did not change before and after heating, and the Ca / P molar ratio calculated from the analytical values of CaO: 55.8% and P ₂ O ₅ : 42.3% was the Ca / P molar ratio of HAP. Of 1.67.

On the other hand, as a high toughness material powder, an average particle diameter of 0.12 μm containing 3 mol% of yttria (Y ₂ O ₃ ) is used.
Was prepared (PSZ-3Y, manufactured by Asahi Kasei Corporation KK), and heat-treated at 1000 ° C. for 2 hours as a pretreatment to remove adsorbed water and volatile impurities.

Next, the spherical HAP particles pretreated as described above and the PSZ powder were mixed in a weight ratio of 10: 0 (comparative example), 10: 2 (example 1), and 10: 5.
(Example 2) and 10:10 (Example 3), and raw material mixtures having four kinds of mixing ratios were prepared.

Next, the raw material mixture for Examples 1 to 3 was
As shown in FIG. 1 and FIG. 2, a high-speed air current impact type powder surface reforming apparatus (Hybridizer, HYB-O type,
K. K. Nara Machinery Works)
The PSZ powder was immobilized (composite) (hereinafter referred to as HYB treatment) so that the surface of the relatively coarse spherical HAP particles was covered with the fine PSZ powder, to prepare composite particles for each example. The processing conditions in the reformer (hybridizer) were such that the rotor rotation speed was 10,000 rpm,
The rotor peripheral speed was set to 60 m / sec, and the processing time was set to 3 minutes.

FIGS. 4 and 5 are electron micrographs showing the particle structure of the composite particles formed by subjecting the raw material mixture in which the mixing ratio of the HAP particles and the PSZ powder was 10: 5 to a composite. It can be seen that the PSZ powder was firmly immobilized so as to completely cover the surface of the coarse spherical HAP particles in multiple layers. Particularly, the average particle size of the PSZ particles is 0.12 μm.
m, the PSZ powder digs into the recesses on the surface of the spherical HAP particles by the impact force during the compounding process by the high-speed airflow impact method, and the bonding strength between the HAP particles and the PSZ powder is sufficiently high by the so-called mechanical anchoring effect. It can easily be inferred that it is compounded with.

At the stage of the composite particles, the bonding strength between the HAP particles and the PSZ powder, which are different materials, has already been sufficiently increased. Particles and PSZ
It is considered that as a result of maintaining the bonding strength with the phase as it is, the high strength characteristics of the entire implant material are further improved.

On the other hand, as a comparative example, four kinds of raw material mixtures in which the mixing ratio of the HAP particles and the PSZ powder were changed as described above were treated in accordance with a normal mixing method, and the corresponding mixed powders for the comparative example were respectively prepared. Prepared. That is, each raw material mixture was mixed by a high-speed stirring type mixer (OMD-3, KK Nara Machinery Co., Ltd., OM dizer) to prepare corresponding mixed powders.

Here, the high-speed stirring type mixer is provided with a stirring blade (blade) rotating at a high speed on the bottom of a cylindrical mixing vessel.
This is a stirring mixer generally used for uniformly dispersing and mixing a plurality of different raw material powders in a short time of about 1 to 3 minutes (hereinafter referred to as OMD treatment).

The processing conditions of the high-speed stirring type mixer for preparing the mixed powder of each comparative example were as follows: the rotation speed of the stirring blade was 500 rpm, the peripheral speed of the stirring blade was 5.4 m / sec.
The processing time was set at 2 minutes.

Next, the composite particles for each example and the mixed powder for each comparative example prepared as described above are sintered (1) after uniaxial pressing (UAP) molding as follows:
(2) Cold isostatic pressing (CIP), followed by sintering, and (3) hot pressing (HP), wherein molding and sintering are carried out simultaneously according to the corresponding examples and comparative examples. Each implant material was manufactured.

That is, in the method of sintering after forming by uniaxial pressing (UAP), first, the composite particles or the mixed powder are mixed with an inner diameter of 1 mm.
The mixture was filled in a 5 mm carbide die and uniaxially pressed at a pressure of 10 to 50 MPa for 3 minutes to form a molded body having a diameter of 15 mm and a height of 10 mm. Then, the obtained molded body was heated to 1000 to 1450 in an argon atmosphere. A process of sintering at 2 ° C. for 2 hours was performed.

In the method of sintering after CIP molding,
First, the composite particles or the mixed powder are preliminarily molded by a handy press to form a columnar molded body having a diameter of 40 mm, and then the columnar molded body is subjected to a handy CIP device (CIP-50-20).
Type 00: Applied Power Japan K. K. Pressure) at a pressure of 203 MPa for 5 minutes to form a molded body.
Sintering was performed at 000 to 1450 ° C. for 2 hours.

In the method of forming and sintering by hot pressing (HP), a hot pressing device (Daia Vacuum Engineering Co., Ltd.
K. 20MP) using composite particles or mixed powder
a. At the same time as pressing with the pressing force of a, heating and baking for 10 minutes at a temperature of 1350 ° C. in an argon gas atmosphere,
An implant material consisting of a 5 × 40 mm sintered body was prepared.

FIG. 6 shows a mixed powder for a comparative example prepared by mixing a spherical HAP particle 30 as a biocompatible particle and a PSZ powder 31 as a high toughness material powder by a conventional stirring and mixing method (OMD treatment). FIG. 3 is a schematic diagram showing a particle structure of an implant material obtained by sintering the mixed powder 40 of the present invention. In the mixed powder 40 obtained by the conventional OMD process, the roughly spherical HAP particles 30
Although dispersed between the SZ powders 31, there are some parts where the HAP particles 30 are directly in direct contact with each other. Therefore, when the mixed powder 40 is molded and sintered to form an implant material, the HAP particles 30 are continuously bonded starting from the directly contacted portion as a starting point, and the HAP continuous phase 50 is easily formed. A continuous phase 5 composed of PSZ as a tough material
1 is also formed, but the fragile HAP having inferior strength characteristics becomes the continuous phase 50, so that the strength varies and the strength characteristics of the implant material as a whole deteriorate.

FIG. 7 shows spherical HAP particles 30 and PSZ.
The particle structure of the composite particles 41 for each example prepared by treating the powder 31 with the method of the present invention, and the composite particles 41
FIG. 2 is a schematic diagram showing a tissue shape of an implant material obtained by sintering a raw material mixture containing the same. In the composite particles 41 for each embodiment formed by the high-speed air current impact method, the spherical HAP particles 30 are completely covered with the PSZ powder 31 and have the covering layer 32. This coating layer 32 is made of HAP
It is firmly fixed to the particles 30 and has a function of isolating adjacent HAP particles 30 so as not to come into direct contact. This isolation effect is exerted even after the composite particles 41 are sintered. As a result, the implant material obtained by sintering is obtained as a three-dimensional structure in which the PSZ phase as a tough material becomes a continuous phase 51.
A structure in which the discontinuous phase 52 is formed such that the HAP particles 30 are independently dispersed in the gaps of the continuous phase 51 is obtained.

As described above, in the implant material according to each example, the PSZ phase having high toughness forms the continuous phase 51, while the relatively fragile HAP particles 30 form the discontinuous phase 52. As a result, the fracture toughness and structural strength of the implant material as a whole are high, and the durability is excellent.

FIG. 8 shows a method of molding and sintering the implant materials according to the comparative example and Examples 1 to 3, and the HAP / P method.
5 is a graph showing a relationship between an SZ weight ratio and a relative density (density ratio) of a material. In the figure, UAP is an implant material produced by sintering a molded body formed by uniaxial pressing at 1450 ° C. (comparative example: 1150 ° C.) for 2 hours, and CIP is a cold isostatic pressing molded body at 1450 ° C. The comparative example is an implant material prepared by sintering at 1150 ° C.) for 2 hours.
Indicates an implant material prepared by sintering at 1350 ° C. for 10 minutes by a hot press method.

As described later, in the implant materials of Examples 1 to 3 manufactured by the uniaxial pressing (UAP) method and the cold isostatic pressure (CIP) method, the zirconia phase change easily occurs during the sintering process. Some of the HAP particles changed to α-type tricalcium phosphate (α-TCP).

With regard to the implant material formed by the uniaxial pressing (UAP) method, the voids between the zirconia powders inside the compact formed of the composite particles are large and the number of point contacts between the powders is small, so that the grain growth is sufficient. It is probable that he did not proceed. Therefore, the relative density of the implant material (sintered body) by the UAP method is 42.8 as shown in FIG.
It remains at a low value of about 51.1%.

In the case of the cold isostatic pressure (CIP) molding method, the relative density of the implant material can be further increased as compared with the case of the UAP method. This is P
It is considered that densification was promoted by increasing the number of point contacts between SZ powders.

Further, in the case of using the hot press (HP) method, the relative density of the implant material can be further greatly increased at any HAP / PSZ weight ratio as compared with the above-mentioned UAP method and CIP method. Was.
This is considered to be because the sintering and densification of the composite particles proceeded more smoothly by the pressing force.

FIGS. 9 to 11 show that the raw material powder having a HAP / PSZ weight ratio of 10/5 is formed into composite particles, and the composite particles are formed and sintered by UAP, CIP, and HP methods, respectively. 9 is a scanning electron micrograph (secondary electron image) showing the particle structure of the fracture surface of the implant material according to Example 2. The breaking of the implant material was performed at room temperature.

As shown in FIG. 9, in the implant material obtained by sintering the HAP / PSZ composite particles after uniaxial pressing (UAP) molding, there are many voids (pores) indicated by black portions and sufficient. It turned out that it was not densified.

On the other hand, as shown in FIG.
In an implant material obtained by sintering a CIP compact of Z composite particles, coarse spherical HAP particles do not come into contact with each other while maintaining the shape and particle size in the raw material state,
The structure which exists in the space | gap part of a continuous PSZ phase in the independent dispersion state is observed. In the PSZ phase, P
It can be seen that the grain growth of the SZ particles has progressed considerably and that voids exist.

Further, in the fracture surface of the implant material, since the spherical HAP particles are cleaved together with the surrounding PSZ structure, the bonding strength between the HAP particles and the PSZ phase is strong. It can be seen that the structural strength of the material increases.

As shown in FIG. 11, HAP / PSZ
Even in an implant material obtained by sintering composite particles by a hot press (HP) method, spherical HAP particles are present in a dispersed state independently of each other in a gap between continuous PSZ phases without contacting each other. Can be confirmed. FIG.
In the above, the black portion is the HAP particles, and the portion observed in a gray network is the PSZ phase.

The implant material obtained by the HP method is C
It is clear that the grain growth of PSZ is less than that of the implant material obtained by sintering after IP molding. The main reasons for this are that the sintering temperature is about 100 ° C. lower than the sintering temperature of the compact by other UAP and CIP methods, and the sintering time is
This is considered to be because it is extremely short at 2. Further, the implant material shown in FIG. 11 has almost no coarse pores and a dense tissue. This dense organization is shown in FIG.
It is clear from the comparison of the relative densities shown in FIG. Also, the fact that the spherical HAP particles were fractured along with the surrounding PSZ structure in the fractured surface confirmed that the joint strength between the two was high.

FIG. 12 shows that the HAP / PSZ weight ratio was 10 /
Example 3 in which the raw material powder of No. 10 was formed into composite particles, and the composite particles were formed and sintered by a hot press (HP) method.
5 is a scanning electron micrograph showing the particle structure of a fractured surface of the implant material according to EDX mapping of constituent elements. The mapping of the constituent elements (Ca, Zr) was performed by using a scanning electron microscope (JSM-JM manufactured by JEOL KK).
5400) with an energy dispersive X-ray spectrometer (JED-2110 manufactured by JEOL KK).

FIG. 13 is a scanning electron micrograph showing the particle structure as a reflected electron image in the same field of view as in FIG. The black granules are HAP particles and the gray mesh is P
SZ phase.

As shown in FIGS. 12 and 13, it can be observed that a continuous PSZ phase in the form of a gray network is formed so as to surround the spherical HAP particles observed as black granules. In particular, the HAP / PSZ weight ratio shown in FIG.
Since the PSZ weight ratio is increased as compared with the / 5 implant material, the network structure of the PSZ phase is more clearly observed as compared with FIG.

Each Example prepared by molding and sintering raw material powders having different HAP / PSZ weight ratios according to the uniaxial pressing (CIP) method, the cold isostatic pressure (CIP) method and the hot pressing (HP) method About the implant material which concerns on a comparative example, it analyzes and investigates a component using a powder X-ray-diffraction apparatus (MXP3V type made by KK MacScience),
The results shown in Table 1 below were obtained. In Table 1, HAP indicates hydroxyapatite, α-TCP indicates α-tricalcium phosphate, ZDO (T) indicates tetragonal zirconia, and ZDO (C) indicates cubic zirconia.

[0091]

[Table 1]

FIG. 14 shows the relationship between the weight ratio of HAP / PSZ of the composite particles used to produce each implant material and the ratio of the weight of HAP in the sintered body (implant material) to the weight of HAP in the composite particles. Shown in

As is clear from the results shown in Table 1 and FIG. 14, the uniaxial pressurization (UAP) method and the cold isostatic pressure (CI
In the implant material obtained by sintering the molded product by the P) method, except for the comparative example in which only HAP was used, the formation of α-TCP in which a part of HAP was changed at any HAP / PSZ weight ratio And a part of ZDO (T) is also changed to ZDO (C). It was confirmed that the amount of α-TCP produced tended to increase with an increase in the amount of PSZ.

In the implant material according to the comparative example composed of HAP alone, it was found that there was no change in the X-ray diffraction pattern before and after sintering, and the proportion of the HAP component was maintained as it was. From this fact, it is considered that the presence of the zirconia component ZDO (T) promoted the change of HAP to α-TCT.

In general, it is known that a considerable amount of HAP particles changes to α-TCP at a sintering temperature near 1200 ° C. even when PSZ powder is not present. The amount of change to is small. The reason is that the HAP particles used in this example show a sharp X-ray diffraction pattern, have extremely high crystallinity, and have a high Ca content.
Since the / P molar ratio is 1.67, which is in agreement with the theoretical value and the purity is extremely high, it is considered that the change to α-TCP did not progress. Therefore, in order to ensure biocompatibility with HAP, it is preferable to use high-purity HAP particles.

On the other hand, in the implant material obtained by molding and sintering by the hot press (HP) method, PS is used.
Even for each composition in which Z exists, α-
No formation of TCP and no phase change of the zirconia component were observed. This is because the sintering temperature is 100 ° C. lower than 1350 ° C. as compared with the UAP method or the CIP method,
It is considered that the heating time was as short as 1/12, that is, 10 minutes.

According to the findings of the present inventors, the particle size is several tens.
When a mixture of fine HAP particles of nm and fine PSZ powder having an average particle diameter of 0.3 μm was sintered at the same processing condition of 1350 ° C. for 10 minutes, a considerable amount of α- The generation of TCP has been confirmed by X-ray diffraction. In this embodiment, the reason why α-TCP is not generated is that coarse HAP particles are used as compared with the conventional example having an average particle size of 9.32 μm, and the particle size ratio between the HAP particles and the PSZ powder is appropriately adjusted. This is probably due to the difference of adjustment.

As is clear from the results shown in Table 1 and FIG. 14, in order to suppress the production of α-TCP in the implant material, increase the residual ratio of HAP, and secure a higher biocompatibility, -It has been confirmed that it is effective to use a hot press (HP) method as a sintering method.

Next, a test piece was cut out from the implant material prepared in each of the examples and the comparative example, and the fracture toughness value was measured. The measurement method is SEVNB (Single) conforming to the three-point bending test method specified in JIS R 1607-90.
Based on the Edge Notched Beam method, the measurement was performed using a desktop precision universal testing machine Autograph (AGS-D type, manufactured by KK Shimadzu Corporation). The test piece size is 3mm × 4
mm × 40 mm, the span was set to 16 mm, and the crosshead speed was set to 0.5 mm / min.

FIG. 15 shows the measurement results of the fracture toughness values. In FIG. 15, the symbol HYB-CIP is a powder surface reforming device (hybridizer) based on the high-speed airflow impact method.
Shows an implant material obtained by molding the composite particles formed by using a cold isostatic pressure (CIP) method and then sintering the composite particles;
Similarly, HYB-HP is capable of hot pressing (H
The implant material obtained by molding and sintering according to the P) method is shown. The symbols OMD-CIP and OMD-HP are implants obtained by molding and sintering a mixed powder prepared using a conventional general-purpose high-speed stirring mixer (OM dither) by the CIP method and the HP method, respectively. Show the material.

As is clear from the results shown in FIG. 15, in the implant material (HYB-CIP) obtained by subjecting the composite particles prepared by the method of the present invention to CIP molding and sintering, the hydroxyapatite (HAP) and the Weight ratio (HAP / PSZ) to stabilized zirconia (PSZ) is 1
No significant change in the fracture toughness value was observed up to 0/2.
6MPam was ^0.5 . In HAP / PSZ weight ratio of 10 / 5-10 / 10 are 0.9MPam ^0.5, this value has increased by 50% compared with the weight ratio of 10 / 0-10 / 2.

On the other hand, in the case of an implant material (OMD-CIP) obtained by subjecting a mixed powder prepared using a conventional high-speed stirring type mixer (OM dicer) to CIP molding and sintering, the HAP / PSZ weight is reduced. When the ratio was 10/5, the fracture toughness increased only to an average of about 0.7 MPam ^0.5 .

On the other hand, the implant material (HYB-HP) obtained by molding and sintering the composite particles prepared by the method of the present invention by a hot press (HP) method is HAP
When the / PSZ weight ratio is 10/2 to 10/5, the fracture toughness value increases to 1.8 MPam ^0.5 and HAP
2.8 MPa when the / PSZ weight ratio is 10/10
m increased to ^0.5 . This value is 1.5 times the fracture toughness value of ordinary crystallized glass for artificial bone, and 30% by volume of PSZ is dispersed in crystallized glass containing HAP and wollastonite. This value is equal to or more than the fracture toughness value of 2.5 MPam ^0.5 of the implant material for living hard tissue.

On the other hand, an implant material obtained by molding and sintering a mixed powder prepared using a conventional mixer (OM diser) having a small impact force acting on the raw material powder by a hot press (HP) method. (OMD-HP), HA
When the P / PSZ weight ratio was 10/5, the average fracture toughness increased only up to 1.3 MPam ^0.5, indicating that the strength properties were low. In addition, it was also confirmed that compared with an implant material (HYB-HP) formed by hot pressing the composite particles, the variation width of the fracture toughness value was large, and a sintered body having stable characteristics could not be obtained.

As described above, among the implant materials according to the present embodiment, the implant material obtained by molding and sintering the composite particles having the HAP / PSZ weight ratio of 10/10 by the hot press (HP) method is 2.8 MPam. It was found to exhibit a particularly high fracture toughness value of ^0.5 . However, the relative density of the implant material (sintered body) is about 78% as shown in FIG. 8, which is still lower than that of other ceramic sintered bodies. Therefore, the pressure during molding, the time and the pressure during sintering,
It has been found that there is room for further improving the fracture toughness value of the implant material by adjusting the processing conditions such as temperature and time and optimizing the relative density with respect to the fracture toughness.

As described above, in the implant material according to the present embodiment, PSZ as a high toughness material exists as a continuous phase having a three-dimensional network structure, and a fragile HAP having biocompatibility is included in the continuous phase. The particles have a microstructure in which the particles are dispersed independently and exist as a discontinuous phase. Thus, the strength properties can be significantly improved compared to conventional implant materials having a relatively fragile HAP particle continuous tissue.

Further, since a region of only the HAP particles having a relatively large area exists on the surface of the implant material,
Since the affinity for bone tissue inherent in the HAP particles is sufficiently ensured, the HAP particles have good fusion with the surrounding living tissue (fitness) after implantation, and are excellent as constituent materials of artificial bones and artificial roots. It is effective.

[0108] In the above manufacturing process of the implant material, when a compact formed by forming the composite particles by the uniaxial pressurization (UAP) method and the cold isostatic pressure (CIP) method is sintered, a part of the HAP component is reduced. α-TCP tends to change, and this change increases with an increase in the amount of PSZ. However, the composite particles are hot-pressed (H
When molded and sintered by the P) method, HAP / PSZ
No formation of α-TCP was observed regardless of the weight ratio. Therefore, in order to particularly suppress the generation of α-TCP and sufficiently secure the biocompatibility with the HAP particles, it is desirable to employ a hot press (HP) method as a molding and sintering method.

Although the present embodiment shows an example in which partially stabilized zirconia (PSZ) is used as the high toughness material, a case where alumina (Al ₂ O ₃ ), which is a ceramic material having little harm to the living body, is used. Also in this example, as in the present example, the bonding strength between the alumina particles and the HAP particles was significantly increased, so that an implant material having excellent strength characteristics was obtained. However, since the toughness value of the alumina sintered body itself was relatively small as compared with the toughness value of the zirconia sintered body, only a strength characteristic lower than that of the implant material of this example using PSZ as a continuous phase was obtained. .

[0110]

As described above, according to the implant material for living hard tissue and the method for producing the same according to the present invention, the high toughness material powder and the particles having biocompatibility are compounded by the high-speed air impact method. Since strong composite particles are used, the two components are less likely to separate again even when pressure-formed or sintered, and the components are less likely to aggregate to form coarse particles. Particularly, particles having biocompatibility are dispersed in a continuous phase composed of a tough material to form a microstructure existing as a discontinuous phase.

Therefore, it is possible to provide an implant material which is excellent in biocompatibility, has high fracture toughness and has affinity for living tissue, and is suitable as a substitute material for living hard tissue such as an artificial bone or an artificial tooth root. it can.

[Brief description of the drawings]

FIG. 1 is a front view showing a configuration example of a powder surface modifying apparatus used for preparing composite particles in the method of the present invention.

FIG. 2 is a side sectional view of a main body of the surface reforming apparatus shown in FIG.

FIG. 3 is a scanning electron micrograph showing the particle structure of spherical HAP particles.

FIG. 4 shows a mixing ratio of HAP particles to PSZ particles of 10: 5.
5 is a scanning electron micrograph showing the particle structure of a composite particle obtained by subjecting a raw material mixture to a composite treatment.

FIG. 5 shows a mixture ratio of HAP particles and PSZ particles of 10: 5.
3 is a scanning electron micrograph showing the particle structure of a composite particle obtained by subjecting the raw material mixture to a composite treatment in a further enlarged manner.

FIG. 6 is a schematic diagram showing a particle structure of a mixed powder prepared by a conventional mixing method and a tissue shape of an implant material obtained by sintering the mixed powder.

FIG. 7 is a schematic diagram showing a particle structure of a composite particle prepared by the method of the present invention and a tissue shape of an implant material obtained by sintering the composite particle.

FIG. 8 is a graph showing a relationship between a molding sintering method, a HAP / PSZ weight ratio, and a relative density of an implant material.

FIG. 9 is a scanning electron micrograph showing the particle structure of a fractured surface of an implant material obtained by molding by a uniaxial pressing method and then sintering.

FIG. 10 is a scanning electron micrograph showing a particle structure of a fracture surface of an implant material obtained by sintering after molding by a cold isostatic method.

FIG. 11 is a scanning electron micrograph showing the particle structure of a fracture surface of an implant material obtained by molding and sintering by a hot press method.

FIG. 12 is a scanning electron micrograph showing the particle structure of the fracture surface of the implant material according to Example 3 by mapping constituent elements.

FIG. 13 is a scanning electron micrograph showing the particle structure as a reflected electron image in the same field of view as in FIG.

FIG. 14 is a graph showing the relationship between the HAP / PSZ weight ratio of composite particles and the ratio of the weight of HAP remaining in the implant material to the weight of HAP in the composite particles.

FIG. 15 is a graph showing a relationship between a HAP / PSZ weight ratio and a fracture toughness value of an implant material.

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 Powder surface reforming apparatus (hybridizer) 2 Main body casing 3 Rear cover 4 Front cover 5 Rotor 6 Impact pin 7 Rotating shaft 8 Impact ring 9 Open / close valve 10 Valve shaft 11 Actuator 12 Circulation circuit 12a Inlet 12b Outlet 13 Raw material hopper 14 Material supply chute 15 Open / close valve 16 Impact chamber 17 Modified powder discharge tube 18 Back collector 19 Preprocessor 20 Material measurement feeder 21 Jacket 30 Biocompatible particles (HAP particles) 31 High toughness material powder (PSZ powder) 32 Coating layer (PSZ coating layer) 40 Mixed powder 41 Composite particle 50 HAP continuous phase 51 PSZ continuous phase 52 HAP discontinuous phase

 ──────────────────────────────────────────────────続 き Continuation of the front page (56) References JP-A-5-220177 (JP, A) JP-A-3-32676 (JP, A) JP-A-55-8797 (JP, A) 504403 (JP, A)

Claims (11)

(57) [Claims] 1. An implant material for a hard tissue of a living body, characterized in that it has a microstructure in which particles having biocompatibility are dispersed in a continuous phase made of a tough material and present as a discontinuous phase. Biological hard tissue characterized in that high-toughness material particles are coated on the surface of each particle having affinity, and the coated bio-affinity particles are dispersed in a continuous phase composed of a high-toughness material. For implant material. 2. The implant material for living hard tissue according to claim 1, wherein the high toughness material is at least one of partially stabilized zirconia and alumina. 3. The implant material for living hard tissue according to claim 1, wherein the high toughness material is at least one metal material selected from stainless steel, titanium, a titanium alloy and a cobalt-chromium alloy. 4. The implant material for living hard tissue according to claim 1, wherein the biocompatible particles are at least one selected from hydroxyapatite, tricalcium phosphate and bioglass. 5. The implant material for living hard tissue according to claim 1, wherein the weight ratio of the high toughness material is 20% or more. 6. The implant material for living hard tissue according to claim 1, wherein the average particle size of the high toughness material particles is 1/5 or less of the average particle size of the biocompatible particles. 7. The implant material according to claim 1, wherein the biocompatible particles are exposed on the surface of the implant material. 8. The implant material for living hard tissue according to claim 1, wherein the implant material for living hard tissue is an artificial bone or an artificial tooth root. 9. A method for producing an implant material for living hard tissue having a fine structure in which particles having biocompatibility are dispersed as a discontinuous phase in a continuous phase made of a high toughness material. The high-toughness material powder is immobilized by a high-speed air current impact method so that the particle surface is coated with the high-toughness material powder to prepare composite particles. A method for producing an implant material for living hard tissue, comprising sintering in an oxidizing atmosphere. 10. The implant material for living body hard tissue according to claim 9, wherein the average particle size of the high toughness material powder is set to 1/5 or less of the average particle size of the particles having biocompatibility. Production method. 11. The method for producing an implant material for living hard tissue according to claim 9, wherein the pressure molding operation and the sintering operation of the composite particles are simultaneously performed by a hot press method.