JP3455382B2 - Implant material - Google Patents

Description

Detailed Description of the Invention

[0001]

TECHNICAL FIELD The present invention relates to an artificial bone material, a bone fixing and joining material or a bone prosthesis material, which has excellent bioactivity by combining with bone tissue, has a high strength, and a partial substitute material for an artificial hip joint. And implant material for living body useful as artificial tooth root, root canal filling material, bone repairing and filling material, artificial tooth material, etc.
About the fee.

[0002]

2. Description of the Related Art In an aging society, femoral fractures and arthritis have become major problems, and the development of artificial hip joints that can withstand long-term use is strongly sought in order to obtain a rich and healthy life. Is desired. In addition, as a prosthetic material for bones and teeth lost due to accidents or diseases, it is strongly desired to develop an implant material that is embedded in the living body and bonds with living tissue. One of them is a composite material composed of metal and ceramic. Is being developed. That is, it is intended to obtain mechanical strength with a metal and biocompatibility with a ceramic.

The inventors of the present invention formed a glass ceramic layer in which calcium phosphate is dispersed in glass as an implant material composed of such a composite material on a metal substrate, and the surface of the glass ceramic layer was dissolved and etched by an acid for countless times. We have proposed that the holes of the
No. 226), and further, by immersing such an implant material in an artificial body fluid, hydroxyapatite excellent in biocompatibility is deposited on the surface, thereby proposing a method for accelerating the bond between the implant material and the bone ( Japanese Patent Laid-Open No. 4-
276257).

[0004]

The main characteristics required for implant materials are mechanical strength, biocompatibility, and safety, and the above-mentioned implant materials proposed by the present inventors sufficiently satisfy these characteristics. In addition, it was also excellent in strength in a state where the implant material and the bone were sufficiently bonded.

However, according to the above-mentioned conventional method, when forming a glass-ceramic layer in which calcium phosphate is dispersed in glass on a metal substrate, it is necessary to continuously coat the dispersed calcium phosphate with a concentration gradient. In order to form a bioactive surface having a structure in which calcium phosphate is exposed by dissolving and etching the surface glass with acid, the properties of calcium phosphate used (for example, chemical stability, solubility, crystallinity, etc.) are required. There was a problem that it greatly depended on the degree of change.

That is, hydroxyapatite [Ca ₁₀ (PO ₄ ) ₆ (OH) ₂ ] (hereinafter referred to as “H
For example, the theoretical density of the crystal is 3.
Was 16g / cm ^3, whereas, taking as an example an alumina borosilicate glass as the glass, its density 2.454
It is g / cm ³ . Therefore, a powder obtained by mixing HA powder and glass powder in a desired ratio (HA / glass weight ratio at this time is 9/1, 7/3, 5/5, 3/7, 1/9
The densities of the dispersed mixtures are 3.072 and 2.90, respectively.
(9, 2.763, 2.630, 2.510 g / cm ³ ) is applied and baked on the base material, and the HA fine particles settle to the bottom due to the melting of the glass, and the concentration gradient is It is difficult to have a continuous coat, and H is also used for dissolution etching of the outermost surface coated glass layer with an acid.
There were severe restrictions, including the nature of A.

[0007]

Means for Solving the Problems The present inventors have solved the problems described above, and have a small composition and structural unevenness, and easily produce a bioactive implant material with good reproducibility even when a commercially available calcium phosphate salt is used. Successful.

That is, the present invention is an implant material, which comprises a glass ceramic layer having calcium phosphate on at least the surface thereof, a glass powder and a calcium phosphate powder.
Mix and disperse, and heat the mixed and dispersed powder to
Caret obtained by melting lath powder and then cooling and solidifying
Calcium phosphate-dispersed cullet fine powder obtained by crushing lumps,
It is characterized in that it is provided by coating on a metal substrate and firing .

[0009]

[0010]

BEST MODE FOR CARRYING OUT THE INVENTION The metal of the metal substrate used in the implant material of the present invention is not particularly limited,
Titanium; Ti-6Al-4V alloy, Ti-6Al-4V
+ 20Vol% Mo, Ti-6Al-4V + 40Vol
% Mo, titanium alloys such as Ti-6Al-7Rh; Ni
-Cr alloys; Co-Cr alloys, stainless steel and the like can be mentioned. Of these, titanium and titanium-based alloys are preferable in terms of excellent in-vivo corrosion resistance and good compatibility with living bodies, and Ti-Al-based alloys are particularly preferable because of their high material strength, and recently, quite complicated shapes. It is possible to perform precision and fine processing even for things.

In the present invention, the glass ceramic layer formed on the metal substrate is not particularly limited in form as long as it has calcium phosphate on the surface, but preferably the content of calcium phosphate in the glass is toward the surface. Is dispersed with a continuously increasing concentration gradient, has a roughness suitable for binding to bone and a myriad of pores on the surface, and the calcium phosphate is exposed. The same applies when the intermediate glass layer is provided.

The calcium phosphate used in the present invention is hydroxyapatite (HA) [Ca ₁₀ (PO ₄ ) ₆
(OH) ₂ ], α-tricalcium phosphate [α-TCP:
Ca ₃ (PO ₄ ) ₂ ], β-tricalcium phosphate [β-T
CP: Ca ₃ (PO ₄ ) ₂ ], octacalcium phosphate [OC
_{P: Ca 8 H 2 (PO} 4) 6 · 5H 2 O ], brucite [B
RUSHITE: CaHPO ₄ .2H ₂ O] and the like, and preferably contains a large amount of HA, and Ca / P (molar ratio) is preferably in the range of 1.50 to 1.75. Incidentally, the Ca / P of HA is 1.67. HA is the main composition of living bone, and the presence of this HA causes the affinity with living bone to be expressed.

Next, as a glass which can be used for the glass ceramic layer or the intermediate glass layer, an alumina borosilicate glass having the following composition is used for bonding strength and linear expansion coefficient with a metal substrate, and further for calcium phosphate during firing. A preferable example is given from the viewpoint that the glass frit hardly reacts.

SiO ₂ + B ₂ O ₃ + Al ₂ O ₃ 75-85 wt% Alkali component 15-20 wt% Here, the proportion of the alkali component is Na ₂ O, K ₂ O, Li.
_It is the total amount of alkali metal oxides such as ₂ O.
If necessary, a small amount of metal oxide such as ZrO ₂ , TiO ₂ and CaF ₂ may be added to the alumina borosilicate glass.

The reason for selecting the compounding ratio of the silica-alumina glass composition is as follows.

If the alkaline component exceeds the above range, the linear expansion coefficient of glass is too large in comparison with the linear expansion coefficient of the metal substrate, and in particular, the firing conditions when firing and manufacturing the implant material according to the present invention are taken into consideration. Then, distortion due to temperature change tends to increase, which is not preferable (by the way,
The linear expansion coefficient of glass is 90 to 9 which is the coefficient of thermal expansion of the metal base material.
It is preferably in the range of 5%. This is based on the fact that glass is strong against compression and weak against tension. ), Adhesion to metal substrates (particularly titanium and titanium alloys) and bond strength tend to decrease. In addition, there is a problem of alkali elution when used as an implant material (the chemical stability and corrosion resistance of glass change with the increase of alkali components), which tends to cause irritation to living tissues and cells. At times, a reaction with a calcium phosphate component occurs, which tends to induce decomposition of calcium phosphate, which is not preferable.

If the amount of the alkali component is less than the above range, the melting temperature of the glass tends to be high, and the coating temperature must be increased. (Alloy) is liable to deteriorate in strength, and further, excessive reaction between calcium phosphate and glass tends to occur, which is not preferable.

The coefficient of linear expansion of glass-ceramics containing calcium phosphate increases with increasing content of calcium phosphate. Therefore, it is possible to control the linear expansion coefficient of the mixture also by adjusting the content of calcium phosphate, and in the implant material of the present invention, particularly in the case of the embodiment in which the intermediate glass layer is interposed, a glass ceramic containing calcium phosphate. The linear expansion coefficient of the glass used for the layer can be any value.

The content of calcium phosphate in the glass ceramic layer is preferably in the range of 15 to 50% by weight when the glass layer is directly provided on the metal substrate.
15-80 when the intermediate glass layer is provided on the metal substrate
It is preferably set to wt%. If the content of the calcium phosphate is less than the above range, biocompatibility tends to deteriorate, which is not preferable. The upper limit of the compounding ratio of the calcium phosphate is 50% by weight when the intermediate glass layer is not interposed. When it exceeds 50% by weight, the bonding strength with the metal base material is lowered and the material strength as an implant is lowered. This is because there is a tendency.

Further, when the intermediate glass layer is provided, the bonding force with the metal base material is increased, and the glass ceramic layer in which calcium phosphate is dispersed and the intermediate glass layer are bonded continuously and firmly and integrally. Therefore, the calcium phosphate content of the glass ceramic layer does not peel itself,
Moreover, it is preferably 80% by weight or less so that elution is not excessive.

Next, a method for producing the implant material of the present invention will be described.

First, regarding the method for producing an implant material in which a glass ceramic layer is directly formed on a metal base material, the glass ceramic layer having roughness and innumerable pores suitable for joining with bone on the surface is taken as an example. explain.

The metal substrate is preferably subjected to degreasing before coating, pickling and blasting. More preferably, the blast treatment is performed so that the average centerline roughness of the metal substrate is 1 to 3.4 μm. Also, after blasting,
You may form an oxide film by heat-processing in 900-950 degreeC temperature under vacuum.

Next, the coating process will be described by taking HA as an example of calcium phosphate.

HA is produced by a known method. Among them, when the wet method is adopted, HA produced by adding H ₃ PO ₄ to Ca (OH) ₂ aqueous solution is dried at 800 ° C. After baking and firing at 1200 ° C., the powder is crushed and adjusted to a predetermined particle size. In the dry method, CaCO ₃ and CaHP
What is synthesized by the sintering reaction of O ₄ (1200 ° C. × 12 hours) or the like is crushed to adjust the particle size. On the other hand, glass is also adjusted to a predetermined particle size.

Next, the glass powder is melted by heating the mixed powder in which HA and the glass powder having the adjusted particle size are mixed and dispersed at a predetermined weight ratio, and then cooled and solidified to form a cullet mass, which is crushed. Prepare cullet microparticles. A flow chart for cullet particle adjustment is shown in FIG.

In producing the cullet lump, the heating temperature is 90
The range of 0 ° C to less than 1100 ° C is preferable, and the range of 900 ° C to 1000 ° C is more preferable. Heating temperature is 11
Above 00 ° C, a slight reaction with molten glass tends to be observed, which is not preferable, and when the heating temperature is below 900 ° C, firing is insufficient (the viscosity of the molten glass phase is high), and the dispersed HA particles become wet. Also, defoaming of bubbles generated tends to be incomplete, and it tends to be difficult to produce a uniform cullet lump (By the way, the softening temperature of glass is 672 ° C, and stirring and melting are recommended when the proportion of glass is large. ).

Further, it is preferable to carry out heating in a vacuum because the glass is melted and trapped between the HA particles and the glass so that bubbles can be easily discharged.

The cullet having the desired composition can be prepared by heating either a combination of cullet powders having different HA contents, a combination of cullet powder and glass, or a combination of cullet powder and HA. According to this method, it is possible to control the generation of bubbles and the strength of the glass ceramic layer. In this case, the heating temperature is lower than the cullet manufacturing temperature above 70.
It is preferably carried out in the range of 0 to 900 ° C.

The heating time is preferably short if a homogeneous cullet mass is obtained, but preferably 4 to 2
It is 0 minutes, more preferably 5 to 15 minutes, and within this time, no reaction between HA and the molten glass phase is observed.

The cullet fine powder is applied and coated on a target substrate (metal substrate, glass layer or glass ceramic layer) and then fired. The coating method is not particularly limited, but it is preferable to perform coating with a concentration gradient so that the HA content continuously increases toward the surface. The concentration gradient may be formed by using cullet fine powder having a constant HA content, or may be formed by providing a plurality of layers using cullet fine powder having different HA contents. When a plurality of layers are provided, the HA content of each layer is preferably within the range of 20-80% by weight.

The firing temperature is preferably in the range of 850 ° C to 1100 ° C. If the temperature is lower than 850 ° C., the firing will be insufficient and the bonding strength with the metal base material tends to be weak. If it exceeds 1100 ° C, the strength of the metal base material (particularly titanium, titanium-based alloy) tends to decrease, and the coexistence of glass tends to cause a decomposition reaction of HA, which is not preferable.

Next, after coating as described above, etching treatment is performed with acid. Etching process is HF and H
It is simple and preferable to use a mixed solution of NO ₃ , but a method of uniformly etching the surface of the specimen in HF vapor for a suitable time is also recommended.

By dissolving and etching the glass with an acid, the surface layer of the glass-ceramic layer has innumerable pores with large irregularities and has a structure in which calcium phosphate is exposed. The size of the holes is preferably several μm to 50 μm.

Regarding the manufacturing method of the implant material in which the glass ceramic layer is provided on the metal base material via the intermediate glass layer, only the step of coating the intermediate glass layer is added, and the others may be the same as above. . The firing temperature at the time of coating the intermediate glass layer is preferably 850 ° C to 1100 ° C.

[0036]

EXAMPLES The present invention will be described in more detail with reference to examples.

Example 1 An alumina borosilicate glass having the composition shown in the table below was used.
It was poured into a metal roll (made of copper) from 400 ° C. and rapidly cooled by splat to prepare a glass frit in the shape of flakes, which was then ground to an average particle size of 20 μm and used as a frit powder.

[0038]

[Table 1] An aqueous solution of high-purity Ca (OH) ₂ (pH 1
2 to 13), an aqueous solution of H ₃ PO ₄ was dropped to obtain a precipitate, which was calcined and fired to synthesize HA.

HA powder pulverized to an average particle size of 20 μm and frit powder 2: 8, 3: 7, 5: 5, 6: 4,
The cullet microparticles were prepared by mixing in a ratio of 7: 3, 8: 2, 9: 1 (HA 20 to 90%) and following the procedure of the flowchart shown in FIG. The cullet particles have a particle size of 200 mesh (74 μm or less).

In FIG. 2, HA and frit were mixed in a weight ratio of 7: 3 (that is, 70 wt% HA, and the following mixing ratio was 70).
% HA) (mixture (a) and its mixture)
Powder X of cullet (b) prepared by heating at 0 ° C for 5 minutes
It is a line diffraction pattern. The diffraction line intensities of HA crystals (the main diffraction planes are shown in parentheses) do not change in both (a) and (b), and the HA crystal fine particles are evenly dispersed in the glass matrix phase even in the cullet and are stable. You can see that it exists.

Next, HA powder (average particle size 20 μm) synthesized by a dry method and commercially available as crystalline HA and glass frit were mixed in various proportions (20 to 90%) in the same manner as above, and the mixture was heated to 900 ° C. , 1000 ℃, 1100 ℃
Was heated at each temperature for 5 to 30 minutes, pulverized and prepared to obtain various cullet fine particles. The reactivity between HA and glass frit under these heating and firing conditions was examined by X-ray diffraction.

3 and 4 show 900 ° C. and 10 ° C., respectively.
00 ℃, 1100 ℃ for 5 minutes, and 900 ℃ for 5 minutes, 1
The result of X-ray diffraction of the cullet powder obtained by heating for 0 minutes and 15 minutes is shown. There is no particular difference except that a slight decrease in diffraction intensity is observed at 1100 ° C. in FIG.
Similar results were obtained with cullet produced using dry or wet synthetic HA powder. Therefore, for the preparation of cullet, a heating temperature of 900 ° C to less than 1100 ° C, preferably 900 ° C to 1000 ° C, and a firing time of 4 to 20 minutes, preferably 5 to 15 minutes are suitable.

FIG. 5 is a scanning electron microscope (hereinafter abbreviated as SEM) photograph of cullet powders having different HA contents prepared under heating conditions of 900 ° C. for 5 minutes. (A), (b) and (c) are 30% HA, 50% HA and 70% respectively
HA. The higher the HA content, the finer the particle size
You can see that it is prepared.

By high-magnification SEM observation, wetting of the crystalline HA fine particles and the glass matrix of the frit melted was extremely good, and the more remarkable it was when heated at a high temperature [By the way, the value of viscosity log η of glass is 4.8 (p at 900 ° C)
oise), 3.3 (poise) at 1100 ° C.], and it was found that the dispersion of HA fine particles and the homogenization of the cullet composition tend to become worse as the heating time becomes longer.

The titanium base material was degreased, pickled, and blasted with an alundum to have an average roughness of 2.3 to 6.7 μm.

30 prepared under heating conditions of 1000 ° C. for 5 minutes
% HA and 40% HA cullet powder were mixed with distilled water 1: 1 respectively and sandblasted Ti and Ti-6
It was spray-coated on an Al-4V substrate (15 × 20 × 0.5 mm), dried at 70 to 90 ° C., and then baked at 900 ° C. for 3 minutes in the atmosphere. Then, the glass film of the outermost surface layer was etched for 2 minutes with a mixed solution of 3% HF and 5% HNO ₃ to obtain an implant material having HA particles exposed on the surface and having numerous pores. Both the 30% HA and 40% HA implant materials have sufficient strength to be sufficiently adhered to the base material and have a structure with severe irregularities.

A cullet powder having an HA content of more than 50% was directly baked and coated on a Ti substrate under the same conditions as above, and the same etching treatment as above was performed, so that more dispersed HA particles were exposed on the surface, In addition, it has a number of pores and forms a preferable surface for joining with uneven bone tissue. To use this implant material in a biomedical implant, the bonding strength between the glass ceramic layer of HA-glass and the Ti substrate is used. There is a tendency for problems to remain. This bonding strength is obtained by applying a Ti base material in a vacuum (10 ⁻¹ to 10 ⁻³ Torr) 8
It is remarkably improved by heat treatment at 50 to 950 ° C. for 5 to 10 minutes.

Example 2 HA powder prepared in the same manner as in Example 1 and HA crystal powder prepared by the drying method were respectively used at 30, 50, and
HA-glass mixtures containing 60, 70, 80 and 90% by weight were prepared, placed in an alumina crucible and a platinum crucible and heated in an electric furnace (in air) at 900 to 1000 ° C. for 5 minutes to prepare a cullet lump.

Those containing 30% and 50% HA include
The molten glass at the time of heating adheres to the inner wall of each crucible used, and the cullet lump cannot be easily taken out, and the yield tends to deteriorate. Therefore, 70% HA content is high
After preparing the cullet of HA, 80% HA and 90% HA, add glass frit to it and heat it at 900 ° C for 5 minutes to obtain HA with 30% or 50% (or any low HA content if necessary). ), A cullet mass was prepared and crushed and prepared to obtain cullet powder.

For HA particles having good crystallinity (whether wet method or dry method), the properties of the cullet prepared as described above (wetting between HA particles and glass or cullet and Ti base material, defoaming and foam structure, Furthermore, it was found that there is no substantial difference in the etching behavior of the HA-glass layer surface obtained by baking-coating cullet powder on Ti and Ti alloy substrates (of course, the shape and particle size of HA particles are different). H due to the difference
Differences are observed in the surface texture such as surface roughness and pore distribution of the A-glass layer, but no difference in the bioactive properties of HA-glass).

Pure Ti rod and Ti-6Al-4V alloy rod (3.1 mmφ × 20 mm, average surface roughness 2.3-)
6.8 μm) for the first layer of glass, the second layer for 30% HA cullet powder (hereinafter simply referred to as 30% HA), and the third layer for 5%.
0% HA and 70% HA are sequentially baked on the 4th layer at 900 ° C. for 2 minutes in the atmosphere (coating with a concentration gradient so that the HA content increases toward the surface)
did. Next, chemical etching was performed to remove the glass coating film of the outermost surface layer to form a bioactive surface.

FIG. 6 shows that a Ti rod is coated with a concentration gradient such that the HA (polycrystalline fine particle) content increases toward the surface as described above (thickness: about 100 μm).
m), HA-glass-Ti whose outermost surface layer is 70% HA
It is a SEM photograph of the implant material surface. (A) is as-coated, (b) is one minute etched with a mixed solution of 3% HF and 5% HNO ₃ , and (c) is a partial enlargement of (b). The surface has a rough structure with severe irregularities and many pores are observed. By etching, bioactive HA particles are dispersed and exposed on the surface to form a structure.

FIG. 7 shows that the HA-glass layer is about 200 times thicker than the case of FIG.
It is a SEM photograph of the Ti implant material surface. (A)
Is coated, (b) is etched with the same mixture for 2 minutes, and (a ') and (b') are (a).
And a partial enlargement of (b).

By etching, the glass matrix of the outermost surface is removed, and the HA particles are exposed to form a surface structure having coarser pores. Furthermore, (a ') shows that the dispersed HA particles have extremely good wettability with the molten glass mother phase, and the glass is continuously and strongly caking fixed to the HA particles, (b').
Clearly show that the dispersed HA particles consist of aggregates of crystalline microcrystals, which are exposed on the surface. These figures suggest that the microstructure of the etched surface of the HA-Glass-Ti implant material is a tissue that is highly suitable for bonding to living hard tissue, especially bone.

Ti base material (15 × 20 × 0.5 mm)
A 70% HA-Glass-Ti implant material was prepared by coating the HA-Glass layer with a concentration gradient and having an intermediate glass layer, as in the case of FIG.
As a result of EM observation (observed by enlarging the cross-section by obliquely polishing the surface at an angle of 5 ° (about 11 times)), the HA fine particles were concentrated in the continuous glass matrix to the surface. Is distributed with a concentration gradient that increases
-It was confirmed that the glass layer has excellent adhesion to the Ti substrate.

In FIG. 8, 80% HA (a) and 90% HA (b) were further coated on the 70% HA-glass-Ti implant material prepared in the same manner as in FIG. 6 (900 ° C.). It is a SEM photograph of the etching surface (1 minute with a mixed solution of 3% HF and 5% HNO ₃ ) of a sample that was baked in the air for 5 minutes. Even when the HA content is extremely high, a relatively uniform coat whose surface is covered with HA fine particles is possible. However, in the case of 90% HA coating, the adhesive force between the HA fine particles exposed on the surface and the glass matrix phase is not sufficient, and it cannot be said that it is in a state suitable for application to a biological implant.

Example 3 In order to evaluate the chemical durability and bioactivity of the HA-glass-Ti implant material, 70% HA-glass-Ti implant material subjected to surface activation treatment (etching treatment) (Fig. 6- ( The electrochemical corrosion behavior of the sample b) similar to b) was measured.

A rod-shaped sample (diameter: 3.2 mmφ × 1) was used for the experiment.
5 mm), an M3 screw thread was provided at one end, and the sample electrode holder was fixed. A platinum electrode was used as the counter electrode, and the anodic polarization curve was measured using the saturated calomel electrode (reference electrode) potential (SC
With reference to E), the corrosion potential was changed to +2 V (volt) at a scanning speed of 1 mV / s, and the current value was recorded on an XY recorder via a logarithmic converter. As the test electrolyte solution, a 0.001N HCl solution (pH = 2.8) and a simulated artificial body fluid (pH = 7.2) were used.

The artificial artificial body fluid has a composition almost equal to the concentration of inorganic salts in human extracellular fluid, and its composition is 137.8.
mM NaCl, 4.2 mM NaHCO ₃ , 3.0 m
M KCl, 1.0 mM K ₂ HPO ₄ , 1.5 mM M
_{gCl 2 · 6H 2 O, CaCl} 2 · 2H 2 O of 2.5 mM,
And 50 mM (CH ₂ OH) ₃ CNH ₂ and 45 mM HCl (pH = 7.1 to 7.4) as a buffer,
Ca / P is 2.5. Table 2 shows the ion concentration of the artificial artificial body fluid.

[0060]

[Table 2] FIG. 9 is an anodic polarization curve of HA-glass-Ti implant materials with different HA contents in 0.001N HCl solution (a) and artificial body fluid (b). In both cases, no pitting potential is observed up to the measurement range of 2V. Also about −
A sharp increase in the current density is seen near the corrosion potential of 0.5 V, and then a slight continuous increase up to 2 V (a state close to saturation). The larger the HA content of the sample, the larger the current density.

As a result of atomic absorption analysis of the test liquid after these measurements, a large amount of Ca was detected. Therefore, it is considered that this high current density is due to the dissolution of Ca from HA. The result of this experiment is the result of the implant material produced by the conventional method (reference: S. Ban, J. Hasegaw).
a and S. Maruno, Biomaterialia
ls, Vol, 12 (1991) P.I. 205), suggesting that the implant material according to the invention exhibits good bioactivity.

Example 4 Ti rods (3.
1 mmφ × 20 mm) with a hole having a diameter of 1.5 mmφ, and using the same method as in Example 2, 70% HA-
A glass-Ti implant material was made. The surface of this implant material was mixed with 3% HF and 5% HNO ₃
Figure 10 after minute etching to form a bioactive surface.
As shown in (3), three to four adult dog femur bones 2 were embedded in one side. After 1 month, 2 months, 3 months,
The implant material 1 of the femur 2 is taken out, and the femur 2
Of the bond strength between the implant material 1 and the implant material 1 with an Instron tester (crosshead pullout continuity 0.5
mm / min) and evaluated. The results are shown in Table 3 in comparison with the conventional product (note that the conventional product does not depend on the coating by cullet preparation, but the HA fine particle powder and the glass frit powder are mixed at a target ratio and directly suspended in distilled water. Application, drying and firing are repeated to form a functionally graded HA-glass-Ti implant material having a HA concentration gradient).

[0063]

[Table 3] After 1 month, the bonding strength tends to be slightly inferior to that of the conventional product, but it can be evaluated that there is no difference considering the standard deviation.
After 2 months, it shows the same and stable strength, and after 3 months, it shows about 20% larger than the conventional product and shows extremely excellent bonding strength with little variation in measured values. From the biohistological observation of this sample, formation of new bone could be clearly confirmed around the implant material.

The bond between the HA-glass-Ti implant material according to the invention and the bone is strong, its pull-out strength is sufficiently large and stable, that the process according to the invention has a concentration gradient consisting of metal-glass ceramics. It can be said that it is suitable for forming a bioactive implant material.

[0065]

As described above, since the implant material of the present invention has a large bond strength with bone and has a good biocompatibility, a cement that is durable and stable and is initially fixed is required. It is expected to be provided to the artificial hip joint system, etc.

That is, in the implant material of the present invention, since the metal base material exhibits strength, the brittleness is eliminated, and the shear strength of the surface layer responsible for bioactivity and the lower glass-ceramic layer thereof is controlled and controlled by air bubbles. Since the tissue having a concentration gradient can be homogenized, the entire complex including the bioactive layer becomes tough.

Calcium phosphate is firmly held by the glass layer, and since the surface layer has innumerable pores and the calcium phosphate is exposed, the presence of the pores and the bioactive substance makes it possible to bond the bone with the living bone. It will be easy. The calcium phosphate having an exposed structure exhibits favorable bioactivity because the elution of its components is suppressed by the stable glass layer firmly adhering.

When the intermediate glass layer is provided, the bonding strength with the metal substrate is further improved, and the content of the calcium phosphate in the dispersed glass layer is continuously varied with a wide range of concentration gradient. It is possible to obtain a biocompatible complex having a wide range of application.

Further , according to the present invention, a glass ceramic layer containing a desired amount of bioactive calcium phosphate can be coated with excellent reproducibility,
Furthermore, the biocompatible composite body can be easily obtained by a treatment operation such as etching with an acid.

Regarding the properties of the calcium phosphate to be used, the high crystallinity required for the properties of the calcium phosphate used for forming a glass ceramic layer of a conventional calcium phosphate-glass-titanium composite and forming a bioactive surface by etching is required. It has the advantage that it can be utilized not only in the dense and dense calcium phosphate crystals, but also in the commercially available spherical, granular or acicular form. Considering that the glass-ceramic layer coating is easy and the reliability of its properties is higher than those of the conventional methods, a great contribution is expected to the practical application of the implant of the present invention to a living body.

[Brief description of drawings]

FIG. 1 is a flowchart of cullet particle adjustment.

FIG. 2 is an X-ray diffraction pattern of a mixed powder of HA and glass and its baked cullet powder.

FIG. 3 is an X-ray diffraction pattern of cullet powder.

FIG. 4 is an X-ray diffraction pattern of cullet powder.

FIG. 5 is a scanning electron micrograph of cullet powder.

FIG. 6 is a scanning electron micrograph of the surface of the implant material.

FIG. 7 is a scanning electron micrograph of the surface of the implant material.

FIG. 8 is a scanning electron micrograph of an etched surface of implant material.

FIG. 9 is an anodic polarization curve of the implant material.

FIG. 10 is an explanatory diagram of a pull-out test method.

[Explanation of symbols]

1 Implant material 2 adult dog femur 3 wires 4 belts

Claims (5)

(57) [Claims] 1. A glass-ceramic layer having calcium phosphate on at least the surface thereof is provided with glass powder and calcium phosphate.
Um powder is mixed and dispersed, and this mixed dispersion powder is heated.
Obtained by melting glass powder by and then cooling and solidifying
Calcium phosphate-dispersed cullet crushed
An implant material, characterized in that a fine powder is coated on a metal substrate and fired . 2. The implant material according to claim 1, wherein the glass-ceramic layer has a rough surface with irregularities, and the calcium phosphate is exposed. 3. The implant material according to claim 1, wherein calcium phosphate is dispersed in the glass ceramic layer with a concentration gradient in which the content continuously increases toward the surface. 4. The implant material according to claim 1, wherein the calcium phosphate contains hydroxyapatite as a main component. 5. The implant material according to claim 1, which has an intermediate glass layer between the glass ceramic layer and the metal substrate.