JPS6346152A - Medical tube and its production - Google Patents

Abstract

(57) [Abstract] This bulletin contains application data before electronic filing, so abstract data is not recorded.

Description

Detailed Description of the Invention [Field of Industrial Application] The present invention relates to a medical tube and a method for manufacturing the same, and more particularly to a medical tube with excellent blood compatibility, particularly a tube suitable for artificial blood vessels, and a manufacturing method thereof. Regarding the method.

[Prior art and its problems] Conventionally, methods for manufacturing medical tubes made of thermoplastic polymer compounds include a method of heating and melting a polymer compound and extruding it (hereinafter referred to as "melt extrusion method"), A method of repeatedly dipping into a polymer solution until a single layer of polymer of the required thickness is formed on the mold (hereinafter referred to as the "dipping method"), and a method of extruding a solution of a polymer compound from an annular orifice and internal solidification from the center of the annular orifice. A method of immersing the entire body in an external coagulating liquid while extruding the liquid (Japanese Patent Application Laid-Open No. 60-18
No. 8164) is known.

However, with these methods, it has not been possible to obtain medical tubes that exhibit blood compatibility with good reproducibility over a long period of time.
In particular, there is no product yet that can be satisfactorily used in artificial blood vessels having an inner diameter of 6 mm or less.

In other words, tubes obtained by melt extrusion have smooth inner and outer surfaces, and the entire wall of the tube has a dense structure that cannot be controlled arbitrarily, and, of course, the needle does not go well when suturing. There is a big practical problem. When such tubes are applied to applications such as artificial blood vessels that require long-term compatibility with blood and surrounding tissue, desirable results can be obtained even when using materials with excellent antithrombotic properties such as polyurethane and polyurethane urea. In other words, if the transplant period is prolonged, calcification will occur over time, resulting in the formation of a large amount of blood clots around the calcification. When used, the intimal tissue extending from the stump of the blood vessel on the living side does not adhere stably to the inner surface of the artificial blood vessel and detaches, causing flow disturbances in that area and forming a thrombus, which can then occur again. It is known that thickening of the intima occurs due to organization.

In the dip method, the dimensional accuracy is low, the wall thickness is uneven, and the wall structure is multilayered, making it impossible to control it with good reproducibility. Since solvent evaporation cannot be uniformly controlled, the structure lacks uniformity, making it impossible to obtain highly reliable products.

The method described in JP-A-60-188164 is characterized in that the solution of the polymeric compound is coagulated from both the inside and outside of the wood, but it is no different from the conventional method for producing hollow fiber membranes. Since solidification proceeds from the inside and outside of the tube, only products with skin layers on both sides can be produced.Also, with this method, it takes a long time for the solidification to proceed to a state where the tube does not deform. When trying to manufacture tubes with large inner diameters that have poor shape stability, deformation and uneven solidification tend to occur during the drawing process, and in reality, if the diameter exceeds 1 to 2 mm, the formed tubes will become distorted. be.

Furthermore, for artificial blood vessels with an inner diameter of 7 mm or less, bondability with biological blood vessels is important, and delicate control of physical properties is essential for improving patency results. In particular, appropriate flexibility is required to smoothly connect the inner surfaces of the living blood vessel and the artificial blood vessel during suturing. In addition, good passage of the suture needle improves suturing performance and has a large effect on patency results. In other words, good anastomotic performance is extremely important for creating a desirable shape of the flow path at the connection with the biological blood vessel. It is. In addition, since artificial blood vessels are used semi-permanently, the wall membrane must withstand the pulsating load of blood pressure over 100,000 times a day, and the seams that serve as stress concentration points gradually expand. The basic property is that it does not tear or tear.

Artificial blood vessels have been used that are made by forming polytetrafluoroethylene into a tube and then stretching it to give it a fine fibrous structure. Although the antithrombotic properties have been improved compared to artificial blood vessels, there are problems such as poor needle passage and bleeding from the needle hole, and there is still room for improvement in connection with biological blood vessels. .

Among artificial blood vessels, there was no one with an inner diameter of 7I1 layer or less, especially one with an inner diameter of 6 m or less, whose patency results were clinically satisfactory and usable. The above-mentioned porous polytetrafluoroethylene is only used for limited purposes, and the patency rate for more than one year is unsatisfactory. The development of blood vessels is desired. In order to improve patency results, it is essential to improve the antithrombotic properties of the material. Furthermore, it must provide the above-mentioned mechanical properties necessary for an artificial blood vessel, and it must maintain a vessel wall structure with excellent long-term patency, and as mentioned above, it must have a dense structure. A large amount of experimental data has shown that artificial blood vessels cannot maintain their patency over a long period of time. In particular, the intima that grows on the inner surface is not stably maintained, and it repeatedly grows and peels off due to blood flow and bending. In particular, abnormal growth of the pannus occurs at the connection with biological blood vessels, and this causes disturbances in blood flow, leading to thrombus growth, which gradually organizes and leads to stenosis of the anastomosis, and stabilization of the intima. For proper engraftment, there is no skin layer on the inner surface and 1 to 1
00, preferably with an indentation with a diameter of 3 to 20 microns, and preferably this indentation penetrates to the vacuole inside the tube wall, so that it does not have a skin layer on at least the inner surface of the tube. An artificial blood vessel having such a structure cannot be made by using a known method of coagulation from both the inside and outside using an annular orifice.

As a method for avoiding this problem, for example, Japanese Patent Application Laid-Open No. 188165/1983 has been proposed. That is, this method prevents the formation of a dense skin layer on the inner surface by mixing a pore-forming agent in a solution, molding it, and then removing it by some method. However, this method not only makes the process extremely complicated, but also involves woody defects such as increased porosity of the blood vessel, i.e., the entire wall of the blood vessel becomes porous, causing plasma exudation and seroma formation. It is a well-known fact that similar complications that worsen the prognosis occur frequently with polytetrafluoroethylene artificial blood vessels, which can lead to infections and necessitate reimplantation. be.

Furthermore, it goes without saying that if a pore-forming agent is used that prevents blood cell components from leaking from the vessel wall, stable tissue engraftment on the inner surface of the artificial blood vessel cannot be expected.

As an artificial blood vessel, it has a hole that penetrates into the inside of the vessel wall to promote stable engraftment of the intima on the inner surface, and a dense structure on the outer surface that prevents not only blood cells but also plasma from passing through. In addition to such a structure, it is desirable to have sufficient mechanical strength, essential antithrombotic properties that can keep the initial thrombus to a small amount, and strong tissue reaction even when transplanted into a living body over a long period of time. It is necessary that the material be made of a material that does not cause deterioration due to biodegradation or cause deterioration due to biodegradation.

In order to eliminate the drawbacks of the conventional medical tubes mentioned above, the inventors of the present invention have conducted intensive research and have succeeded in manufacturing medical tubes with excellent blood compatibility, especially artificial blood vessels with excellent patency. This was a success and led to the completion of the present invention.

[Structure of the Invention] The medical tube of the present invention is a single-layer medical tube made of a polymer compound, and the internal structure of the tube wall constituting the tube is porous, and the inner surface of the tube wall is porous. is characterized by the absence of a skin layer.

As one embodiment of the medical tube of the present invention described above, the internal structure of the tube wall is configured such that a large number of vacuole walls constituting a vacuole are connected in series, and the inner surface of the tube wall is connected to the vacuole wall. An example of this is something that is made up of a series of . The vacuole wall is preferably porous and has a large number of pores that are sufficiently smaller than the vacuole. In the medical tube of the invention, the vacuole wall constituting the vacuole has a micro-sponge-like structure in which fine pores of 0.0 to 30 mm are present, so cells can easily grow there and the healing effect is high.

The medical tube of the present invention has no special internal structure or foreign skin structure, so it has excellent compliance (
mechanical adaptability), excellent biocompatibility, and stabilization of the growing endothelium. In this way, the present invention maintains the original mechanical properties of the material used, such as strength and fatigue resistance, does not undergo plastic deformation over time after implantation, and has a dense internal tissue. This is the first time that we have been able to provide a medical tube that does not have the rigidity of medical tubes and has excellent mechanical properties, compliance, and cell production stability.

Among the medical tubes of the present invention, in the case where a large number of vacuole walls constituting vacuoles are continuously connected throughout the internal tissue of the tube wall constituting the tube, the vacuoles The size of is the diameter of the largest vacuole in any cross section (9,)
is preferably in the relationship of 0.005 d≦min≦0.9 d, and more preferably in the relationship of 0.01 d≦cross≦0.8 d with respect to the thickness (d) of the tube wall. be. Note that if the length exceeds 0.9 d, the mechanical strength is low from a practical standpoint, causing concerns about clinical use, and if the length is less than 0.005 d, preferred compliance cannot be obtained.

Another feature of the medical tube of the present invention is that it has extremely good suturing properties.The quality of the 0M coaptation often affects the patency of the transplanted blood vessel, and poor suturing conditions can cause disturbances in blood flow there. It triggers thrombus formation and thus results in occlusion of the graft vessel.

As in the present invention, by constructing the tube from a polymer compound and making the internal tissue of the tube wall porous, the combination of appropriate elongation and flexibility of the tissue makes it easy for the suture needle to pass through and suture, making it smooth. can be anastomosed to host blood vessels.

It is interesting to note that artificial blood vessels whose internal tissue is porous do not tear from the sutures, but do not tear from the sutures like conventional polyurethane artificial blood vessels, which have a densified part in the internal tissue. The drawbacks caused by this have been completely eliminated. It is also an interesting phenomenon that blood does not leak from the sutured seams, and this is due to the inherent elasticity of polyurethane, as well as the effect of the many layers of the vacuole walls that make up the vacuole. This seems to be due to

When the medical tube of the present invention is anastomosed to a host blood vessel, the internal tissue of both becomes a continuous vacuolar tissue, making it easy for cells to invade from the anastomosis, resulting in excellent healing of the occlusal area. In addition, since there is no skin layer on the inner surface of the tube wall that constitutes the blood vessel, it is suitable for the growth of endothelial cells and becomes biogenic quickly, which means that the medical tube of the present invention has extremely excellent long-term patency. This seems to be the cause.

In the distribution of vacuoles in the cross section of the tube wall constituting the medical tube of the present invention, if the size of the vacuoles is configured to be smaller as it approaches the outer surface, blood leakage from the sutured portion can be reduced. Surgical results can be significantly improved. This is achieved by allowing coagulation to occur only from the outside during manufacture of the medical tube of the invention.

The polymer compound used in the present invention is a substance that is highly compatible with blood and tissues, that is, it does not have acute or chronic toxicity, pyrogenicity, or hemolysis, and does not cause inflammation in surrounding tissues even after long-term implantation. It is a polymer that does not cause Examples of such polymers include polyvinyl halides, polystyrene and its derivatives, polyolefin polymers, polyester condensates, cellulose polymers, polyurethane polymers,
Examples include polysulfone resins and polyamide polymers. Of course, copolymers or mixtures containing these materials may also be used; polyurethane-based materials are preferred from the viewpoint of mechanical properties, in-vivo stability, and antithrombotic properties. .

Specific examples thereof include polyurethane, polyurethane urea, blends of these with silicone polymers, and those having an interpenetrating network structure. These also include segmented polyurethanes or polyurethane ureas, those containing polydimethylsiloxane in the main chain, and those containing fluorine in the hard and soft segments. Polyether-type polyurethane or polyurethane urea is preferable to polyester-type because it is less susceptible to biodegradation.

The polyether constituting the polyether segment such as polyurethane is most preferably polytetramethylene oxide, but other polyalkylene oxides (however, alkylene has 2 and/or 3 carbon atoms) are also preferred.
Specific examples of such polyalkylene oxides include polyethylene oxide, polypropylene oxide, ethylene oxide-propylene oxide copolymers, or block copolymers. Furthermore, a polyurethane having both wood-philicity and mechanical properties that contains a polytetramethylene oxide segment and a polyalkylene oxide (alkylene has 2 and/or 3 carbon atoms) in its main chain may also be used. This polyurethane has outstanding antithrombotic properties and biocompatibility, and is a new type of polyurethane with good biocompatibility discovered by the present inventors.

The molecular weight of the polyether forming these soft segments is usually in the range of 400 to 3,000, preferably 450 to 2,500, more preferably 500 to 2,500.
00, and the best polyether segments have molecular weights of SOO to 2.500, especially molecular weights of 1.3
00 to 2.000 polytetramethylene oxide chains. The molecular weight of this polyether soft segment is 3,
If it exceeds 000, the mechanical properties of the polyurethane artificial blood vessel will be poor, and if it is less than 400, it will be too hard to be used even if it is molded as an artificial blood vessel.

Synthesis of polyurethane is carried out by reacting the above-mentioned polyether with hydroxyl groups at both terminals with diisocyanates used in known polyurethane synthesis, such as 4,4'-diphenylmethane diisocyanate, toluidine diisocyanate, 4,4'-dicyclohexylmethane diisocyanate, and hexamethylene diisocyanate. It can also be synthesized using the conventional method of creating a prepolymer of terminal incyanate and extending the chain with diamines such as ethylene diamine, propylene diamine, and tetramethylene diamine, or diols such as ethylene glycol, propylene gelicol, and butadiol. good.

The medical tube of the present invention has a porous inner tissue of the tube wall constituting the tube, and there is no skin layer on the inner surface of the tube wall, so in addition to the inherent biocompatibility, the tissue is It is soft and therefore extremely easy to suture;
No pannus or suction occurs at the anastomotic site, and it has excellent compliance, so when used as an artificial blood vessel, it moderately elastically deforms with the heartbeat, alleviating the stimulation of blood against the host blood vessel. Furthermore, since there is no skin layer on the inner surface of this blood vessel, cells can easily invade into the vacuole, so it has excellent healing properties, and this is the path to an artificial blood vessel with an inner diameter of 6 m or less and excellent long-term patency. It is opened.

The medical tube of the present invention can be manufactured, for example, as follows.

That is, by extruding a rigid core rod with a circular cross section from a circular orifice, a solution of a polymer compound is forced out through a gap slit between the orifice and the core rod so as to be spread over the entire circumferential surface of the core rod, The core rod can be produced by introducing the core rod into a coagulation bath, coagulating the polymer compound around the core rod, and then taking out the core rod.

It is preferable to set the solution used for molding so that the viscosity at the molding temperature is 5 poise or more. If the viscosity is less than 5 poise, huge voids will be generated inside the tube wall and the strength will deteriorate. descend. Also, unevenness in wall thickness tends to occur during the molding process. If it is 10 poise or more, there are fewer restrictions on molding conditions, so it is more preferable.

On the other hand, there are almost no restrictions on the high viscosity side, and sufficient molding is possible even without fluidity of the solution. When molding using a known method for manufacturing a hollow mold using an annular nozzle, it is difficult to form a hollow mould.
A major feature of the manufacturing method of the present invention is that even a solution of approximately 0 poise can be molded very easily. However, since it is desirable from the viewpoint of production that the solution can be defoamed relatively easily, it is preferably 3000 poise or less, more preferably 2000 poise or less.

In the manufacturing method of the present invention, the solvent used for the solution of the polymer compound can be easily selected from known solvents for each substance, but in order to avoid residue in the product and to reduce the cost of the process. From this point of view, water-soluble solvents are advantageous. Such solvents include, for example, dimethylformamide,
Examples include dimethylacetamide, dimethylsulfoxide, N-methyl-2-pyrrolidone, dioxane, tetrahydrofuran, acetone, and the like. Further, in the production method of the present invention, the solution does not necessarily have to be in a good dissolved state, and therefore a large amount of a poor solvent or a swelling agent such as urea can be mixed and used. This is extremely advantageous for the manufacture of medical tubes, particularly artificial blood vessels, which is the object of the present invention. That is, by selecting a wide range of solvent systems, it is possible to easily change the porosity over a wide range without particularly requiring complicated steps such as using a pore-forming agent.

In the manufacturing method of the present invention, the rigid rod used as the core is made of a material that does not dissolve in the solution and does not easily change shape until it is introduced into the coagulation bath. Corrosion resistance is also required, so stainless steel, steel, or brass coated with chrome or Teflon is particularly preferred.

The extruded core rod with the solution spread over its entire circumferential surface is led to a coagulation bath either directly or after passing through a certain dry section.

In other words, there are two methods: wet coagulation, in which a polymer compound solution discharged from a slit between a circular orifice and a core rod is directly discharged into an aqueous coagulation bath, and dry-wet coagulation, in which it is introduced into an aqueous coagulation bath after passing through a dry section. Either is applicable.

When applying the known hollow fiber membrane spinning method to something with a large diameter and wall thickness, such as an artificial blood vessel, it is difficult to maintain the desired shape unless rapid coagulation is applied from both the inside and outside. be. Therefore, conditions that exert a strong coagulation effect from both the inner and outer surfaces are essential for the stable manufacture of tubes, but this is a major obstacle in creating a desirable structure for an artificial blood vessel. That is, a denser structure is formed on the inner and outer surfaces than on the inside of the tube wall.

These drawbacks have been solved by the manufacturing method of the present invention.

That is, since the rigid body existing inside plays the role of stably holding the surrounding solution, it has become possible to produce tubes with uniform shape and size with good reproducibility only by solidifying from the outside surface.

Furthermore, even if the rate of solidification from the outside is slowed down, there is no problem in molding, and solidification can take a sufficient amount of time.

In the method of the present invention, since the solidification conditions can be varied over a wide range, there is a large degree of freedom in molding, and molded products with various structures can be produced.

By adding, for example, a solvent, a poor bath agent, a salt, etc. to the coagulation bath, the structure of both the inner and outer surfaces or the outer surface can be changed widely.

In addition, the surface energy of the rigid body used as the core can be changed by material selection, Teflon, silicone coating, etc., and not only the microscopic properties that affect antithrombotic properties, but also the morphological surface condition can be changed. It can be changed.

One of the distinguishing features of the production method of the present invention is that the passage time through the dry section can be very precisely controlled when wet-dry coagulation is carried out. In particular, the passage time of this dry section is 0.01~
Accurate control is possible even in a short period of several tens of seconds. This is extremely important in making the outer wall structure of the formed tube uniform, and is significantly different from the conventional dip method.

For the above reasons, it is preferable to proactively provide a dry section. In this case, a smooth outer surface can be easily obtained by utilizing the flow of the solution of the polymer compound on the surface in the dry section. The optimal length of the dry zone is determined mainly by the viscosity of the solution and the volatility of the solvent, but is usually 5 to 30 mm.
A range of 0 m+s is desirable; if it is less than 5 mm, the above-mentioned surface smoothing effect will be insufficient. If it exceeds 300 mm, the solution will flow down, resulting in uneven thickness or a product whose thickness gradually changes in the length direction. Furthermore, when a highly volatile solvent is contained in a solution of a polymer compound, evaporation progresses from the surface, and as a result, the surface may reach the dew point, resulting in phenomena such as minute water droplets condensing. , requiring strict control of environmental conditions.

For the above reasons, the length of the dry part is preferably 20 to 20
0m is within the proverbial range.

The extrusion speed when the core rod passes through the dry section is usually 1~
300+sm/sec, preferably 5-200mm/sec
seconds, more preferably in the range of 10-100 mm/second. When the extrusion speed exceeds 300+am/sec, distortion due to residual stress tends to appear in the post-processing process, and 1 ra
If it is less than ts/second, the structure of the outer surface may vary depending on the temperature and humidity of the atmosphere and the concentration of the solvent.

For the coagulation bath, in order to remove the solvent, it is preferable to use one with excellent compatibility.Water is particularly preferable from the viewpoint of safety and cost.If necessary, methanol, ethanol, imbropanol denatured bath may be used. Lower alcohols such as alcohol may be used.In any case, in order to completely remove the solvent, it is preferable to use a solvent that can ultimately replace water in the coagulation system1.For example, if water is added to the coagulation bath When used, a structure that generally becomes sparser from the outer surface to the inner surface is obtained, and since coagulation from the inside does not act at all on the inner surface, the inner membrane does not form a foreign dense layer. The above-mentioned rough surface structure of about 1 to 100 is required for the film to be kept thin and stable. This depression penetrates into the internal cavity and contributes to the stabilization of the intima.

The obtained coagulated molded product may be sterilized after sufficiently removing the solvent, air-dried or forced-dried, or replaced with physiological saline while still wet, and sterilized by autoclaving or gamma rays.

When using polyurethane or polyurethane urea,
A porous structure can provide a preferable compliance (C), which can be expressed as follows: % formula % () (where vo is the internal volume of the measured blood vessel when the internal pressure is 50 mmHg, and ΔP is the internal volume from the internal pressure of 50 mmHg. 150mm) I
Change amount when changing to g10C)++aHg, Δv
is the internal volume of the measured blood vessel that increased when the internal pressure changed by 50 ml II Hg. ) If the value exceeds 65%, it will cause gradual irreversible expansion of the lumen due to repeated pressure stress after long-term implantation, and if this value is too low, In particular, if it is less than 1%, it causes abnormal expansion on the host blood vessel side, resulting in turbulent flow near the anastomosis, which causes thickening of the intima of the proboscis.

Therefore, it is preferable to determine the extrusion thickness and the polymer concentration of the solution in advance so that the compliance is usually 1 to 65%, preferably 3 to 20%, and more preferably 3 to 10%.

Alternatively, the method of the present invention may be carried out after covering the core rod with a reinforcing material such as a polyester mesh and embedding the reinforcing material inside. Alternatively, the core rod may be thinly coated with another material in advance. After that, the method of the present invention may be carried out.

[Examples of the Invention] Hereinafter, the present invention will be explained in more detail with reference to Examples, but these Examples are not intended to limit the scope of the present invention in any way.

Note that all (%) shown below indicate weight % unless otherwise specified.

Support 1 Polytetramethylene glycol with a molecular weight of 1650 and having hydroxyl groups at both ends was reacted with 4,4'-diphenylmethane diisocyanate to produce a prepolymer of isocyanates at both ends, and this was chain-extended with butanediol to synthesize polyurethane. The polyurethane was purified by reprecipitation with a tetrahydrofuran-ethanol system and refluxing three times. This purified polyurethane was dissolved in dimethylacetamide and 1
It was made into an 8.3% solution.

From a circular orifice with a diameter of 10 mm, a stainless steel rod with an outer diameter of 8Il111, which is precisely set to be concentric with the orifice, is extruded at a constant speed, and the entire length of the rod is extruded from a uniform gap between the extruded stainless steel rod and the orifice. The solution was uniformly extruded so that the polyurethane droplets were cast on the circumferential surface, and the rod was introduced into water at 30° C. and solidified.

In this case, since the core is a stainless steel rod, solidification occurs only from the outside. After 3 hours, the rod was taken out and soaked in running water overnight to thoroughly remove the solvent.

The thus solidified polyurethane tube was peeled off from the stainless steel rod and air-dried.

The obtained polyurethane tube had a milky white porous structure, and the cross section of the tube wall was in a state in which a large number of vacuole walls constituting a vacuole were connected in succession. The thickness of the tube wall is 0.7
+u+, and the inner diameter of the tube wall was 7 cm due to the shrinkage during air drying. There is no foreign skin layer on the inner surface of the tube wall of the polyurethane artificial blood vessel obtained in this way,
In addition, vacuoles are not opened and directly exposed on the inner surface of the tube wall, so the inner surface of the tube wall has microscopic pores similar to those of the vacuole wall in the internal tissue of this artificial blood vessel. It had a sponge structure. When observed under 200x magnification with a scanning electron microscope, it was found that the physical structure of the inner surface of the vessel wall was essentially the same as that of the vacuole wall in the internal tissue of this artificial blood vessel, and that there was a gap between the two. No difference was observed. The inner surface of the wall of this artificial blood vessel had a structure that did not cause swirling of blood flow, which would occur if the vacuole were directly exposed by opening into the inner surface of the wall. In addition, the vacuoles were made up of large bubbles, and were relatively small near the outer surface of the tube wall. Moreover, the compliance value was 45%. Fill the polyurethane tube obtained in this example with water,
This water was kept at a positive pressure of 120 + * + * Hg and was used for extracorporeal blood circulation in artificial dialysis (needle diameter 1, 5 m).
m) was repeatedly punctured at random positions over a length of 15c1 of the tube 150 times to check for water leakage. Despite 150 punctures, no significant water leakage was observed.

1凰U An experiment completely similar to Example 1 was conducted using a stainless steel rod with an outer diameter of 4 ms in this example. As a result, the inner diameter of the tube wall is 3.
A compliant tube (C = 64%) was obtained with a tube wall thickness of 2 mm and a tube wall thickness of 0.5 mm, in which the internal tissue of the tube wall was composed of a large number of continuously connected vacuole walls. .

The inner surface of the tube wall of the tube according to this example is composed of the vacuole walls connected continuously, as in the case of the tube according to Example 1, and there is no skin layer (more dense layer) on this inner surface. It didn't exist.

This tube was used as an artificial blood vessel and was implanted into the fibular artery and femoral artery of an adult mongrel dog by suturing one end to the other. The anastomosis was extremely easy with moderate flexibility and compliance (39%), and there was no bleeding from the suture site. Six animal experiments were conducted under the same conditions, but each transplanted blood vessel was
It was still patent 6 months later.

Polytetramethylene glycol with a molecular weight of 1650 and having hydroxyl groups at both ends was reacted with 4,4'-diphenylmethane diisocyanate to form a prepolymer with incyanate at both ends, and this was chain-extended with butanediol to synthesize polyurethane. The synthesized polyurethane was purified by repeating reprecipitation three times in a tetrahydrofuran-ethanol system. This purified polyurethane was dissolved in dimethylacetamide and 2
A 0.0% solution was prepared.

A stainless steel rod with an outer diameter of 7ts, which is precisely set to be concentric with the orifice, is extruded from a circular orifice with a diameter of 10 mm at a constant speed, and the entire length of the rod is extruded from a uniform gap between the extruded stainless steel rod and the orifice. The polyurethane solution was uniformly extruded so as to be spread over the circumferential surface, and the rod was introduced into water at 20° C. and solidified.

In this case, since the core is a stainless steel rod, solidification occurs only from the outside. After 10 hours, the rod was taken out and soaked in running water overnight to thoroughly remove the solvent.

The thus solidified polyurethane tube was peeled off from the stainless steel rod and air-dried.

The obtained polyurethane tube had a milky white porous structure, and the cross section of the tube wall was in a state in which a large number of vacuole walls constituting a vacuole were connected in succession. The thickness of the tube wall is 1.1
mm, and the inner diameter of the tube wall was 6.5 mm due to some shrinkage during air drying.

There is no foreign skin layer on the inner surface of the tube wall of the polyurethane artificial blood vessel obtained in this way, and no vacuoles are opened and directly exposed on the inner surface of the tube wall. The inner surface of the wall had a microscopic spongy structure with fine pores similar to those of the vacuole wall in the internal tissue of this artificial blood vessel. When observed under 200x magnification with a scanning electron microscope, it was found that the physical structure of the inner surface of the vessel wall was essentially the same as that of the vacuole wall in the internal tissue of this artificial blood vessel, and that there was a gap between the two. No difference was observed. The inner surface of the wall of this artificial blood vessel had a structure that did not cause swirling of blood flow, which would occur if the vacuole were directly exposed by opening into the inner surface of the wall. In addition, the vacuoles were made up of large bubbles, and were relatively small near the outer surface of the tube wall. Moreover, the compliance value was 25%. The polyurethane tube obtained in this example was filled with water, and the water was
Although the wing was maintained at positive pressure, no water leakage from the pipe wall was observed.

Mitate 1 Exactly the same experiment as in Example 3 was carried out using an outer diameter of 4.5 mm in this example.
This was done using a stainless steel rod. As a result, the inner diameter of the tube wall is 4.0-■, the thickness of the tube wall is 0.5 mm, and the internal structure of the tube wall is constructed in a state where many vacuole walls constituting the vacuole are connected in a continuous manner. A tube with C=19%) was obtained.

The inner surface of the tube wall of the tube according to this example is composed of the continuous vacuole walls, as in the case of the tube according to Example 3, and there is no skin layer (more dense layer) on this inner surface. It didn't exist.

This tube was used as an artificial blood vessel and was implanted into the iliac artery and femoral artery of an adult mongrel dog by suturing one end to the other.

Compared to commercially available Teflon artificial blood vessels, the needle fit and adhesion to the host blood vessel were better, and no bleeding from the needle hole or anastomotic site was observed after surgery under 1000 units of systemic heparinization. In an experiment using 3 cases, good patency was shown on the gold side after 18 months.

Note 2 "T-" Using the polyurethane solution used in Example 1, apply it to an annular nozzle using a gear pump (the solution outlet has an inner diameter of 31 mm).
The solution was extruded at a rate of about 40 c+s/min from a solution with an outer diameter of 11 m and an outer diameter of 5 m11, and at the same time, pre-defoamed water was discharged from the center inside the solution being extruded in an annular shape. The amount of water discharged was 1.2 times the amount of extruded polyurethane solution. The extruded linear body containing water was immediately introduced into water to solidify the polyurethane into a tubular shape, and was left as it was for an additional hour.

The tube thus obtained had an inner diameter of about 3 mm,
Since the outer diameter was approximately 4.2 m + s and it solidified from both the inside and outside, there existed a skin layer on both the inner and outer surfaces of the tube wall that was different from the internal structure of the tube wall and was denser than this. was. This is a distinctive feature that is different from the tube obtained in Example 1.2, which has already been described. Was. The tube of this reference example had a sponge-like structure in the cross section of the tube wall, which was very similar to that of Example 2. Using the tube of this reference example, a mongrel adult dog was tested under exactly the same conditions as Example 2. was transplanted to. Three of the four transplanted blood vessels were occluded after one month and one after three months.

The dog was sacrificed and a kidney autopsy revealed that the capillaries in the kidney were blocked. When a similar autopsy was performed on the dog of Example 2, no abnormality was found in the kidneys.

Reference 2" 1-e Using the polyurethane solution used in Example 3, apply it to an annular nozzle using a gear pump (the solution outlet has an inner diameter of 3").
water was extruded from the inside center of the extruded solution in an annular shape. The amount of water discharged was 1.2 times the amount of extruded polyurethane solution. The extruded linear body containing water was immediately introduced into water to solidify the polyurethane into a tubular shape, and was left as it was for an additional hour.

The tube thus obtained has an inner diameter of approximately 3.3I
, the outer diameter was approximately 4.81 mm, and it solidified from both the inside and outside, so there was a skin layer on both the inner and outer surfaces of the tube wall that was different from and denser than the internal structure of the tube wall. was. This is a distinctive feature that is different from the tube obtained in Example 3.4, which has been previously described. had. In the tube of this reference example, the cross-sectional appearance of the tube wall exhibited a sponge-like structure very similar to that of Example 4. When the cross section of the tube was examined, it was found that it was flattened and the two orthogonal inner diameters X and Y had a relationship of X/Y = 1.30. Using the tube of the reference example, it was exactly the same as Example 4. Transplantation was performed in adult mongrel dogs under certain conditions. One of the two transplanted blood vessels was occluded after 3 months, and the other patient was sacrificed after 2 months and ri! Upon examination, a pannus of about 1 cm was observed at the anastomotic site, and a large amount of new thrombus had formed around it.

A prepolymer with incyanate groups at both ends is prepared by a conventional method from polytetramethylene glycol having hydroxyl groups at both ends of Mitate U molecule 11890 and 4,4'-dicyclohexylmethane diisocyanate, and this is chain-extended with ethylenediamine to form polyurethane. Synthesized urea. This was purified by reprecipitation three times using a dimethylformamide-ethanol system.

This polyurethane urea was dissolved in dimethylformamide to form a solution with a concentration of 20%. A stainless steel rod with an outer diameter of 5 mm is pushed concentrically from inside a circular orifice with a diameter of 7.2 m, and the polyurethane urea is applied to the entire outer periphery of the rod from a uniform gap between the extruded stainless steel rod and the orifice. The solution was uniformly extruded so as to be cast, and the outflow rate of the polyurethane solution was matched with the extrusion rate of the stainless steel rod, and then introduced into water at 10°C. The extruded polyurethane tube was slowly solidified from the outside, and a white polyurethane film was formed around the stainless steel rod after about 30 minutes. This was left to stand day and night to complete coagulation, and was further immersed in clear water for 20 hours to completely remove dimethylformamide. The obtained polyurethane urea tube was peeled off from the stainless steel rod and air-dried at room temperature.

The inner diameter of this polyurethane urea tube after drying is 3.
6mm, the thickness of the tube wall is 0.6I+m, and the cross-sectional structure of this tube wall is essentially a structure in which many connected vacuole walls, which constitute a vacuole, are formed in its entirety, except for a thin skin layer on the outer surface of the tube wall. I took it. That is, as in the case of the tube according to Example 1, the cross-sectional structure of the tube wall is such that, except for the skin layer, a large number of vacuole walls constituting the vacuole are continuously connected, and the inner surface thereof is It consisted of the vacuole walls connected continuously. Further, the compliance value of the artificial blood vessel of this example was 62%.

This tube had appropriate softness, elasticity, and compliance and was easy to handle. Using this tube as an artificial blood vessel, arteriovenous bypass surgery was performed in which the tube was implanted as a femoral artery-femoral vein bypass in an adult mongrel dog. The joining method was end-side joining, and the total length of the bypass was 20c layers. The N-synthesis was excellent, the needle was easy to pass through, and no blood leakage was observed after suturing, and the bypass tube remained patent even after 6 months.

A prepolymer having incyanate groups at both ends was prepared by a conventional method from polytetramethylene glycol having a molecular weight of 1890 and having hydroxyl groups at both ends and 4,4'-dicyclohexylmethane diisocyanate, and this was chain-extended with ethylenediamine. , synthesized polyurethane urea. This was purified by reprecipitation three times using a dimethylformamide-ethanol system.

This polyurethane urea was dissolved in dimethylformamide to make a solution with a concentration of 33% (viscosity 4300 poise at 20°C), and a stainless steel rod with an outer diameter of 5 m and a diameter of 8.0
Push concentrically from inside a circular orifice of mm in diameter, and extrude the solution uniformly so that the polyurethane urea solution is cast around the entire outer circumference of the extruded stainless steel rod through a uniform gap between the extruded stainless steel rod and the orifice. The outflow rate of the polyurethane solution was matched with the extrusion rate of the stainless steel rod, and the solution was introduced into water at 20°C. The extruded polyurethane tube was slowly solidified from the outside, and a white polyurethane film was formed around the stainless steel rod after about 30 minutes. This was left to stand day and night to complete coagulation, and was further immersed in running water for 20 hours to completely remove dimethylformamide. The obtained polyurethane urea tube was peeled off from the stainless steel rod and air-dried at room temperature.

The inner diameter of this polyurethane urea tube after drying is 4.
The thickness of the Shikokugumi tube wall is 0.6 mm, and the cross-sectional structure of this tube wall is essentially a structure in which many vacuole walls constituting a vacuole are connected, except for a thin skin layer on the outer surface of the tube wall. I was on target. That is, as in the case of the tube according to Example 3, the cross-sectional structure of the tube wall is configured such that a large number of vacuole walls constituting the vacuole are connected in series, excluding the skin layer,
Its inner surface consisted of a series of connected vacuole walls. Further, the compliance value of the artificial blood vessel of this example was 25%.

This tube had appropriate softness, elasticity, and compliance and was easy to handle. Using this tube as an artificial blood vessel, arteriovenous bypass surgery was performed in which the tube was implanted as a femoral artery-femoral vein bypass in an adult mongrel dog. The joining method was end-side joining, and the total length of the bypass was 20 cm. The N-synthesis was excellent, the needle was easy to pass through, and no blood leakage was observed after suturing, and the bypass tube remained patent even after 6 months.

1. Into the polyurethane and polyurethane urea solutions used in Examples 1 and 5, stainless steel rods with an outer diameter of 4 mm and 5 mm were added using a conventionally known method, that is, dip drying was repeated 10 times or more. , polyurethane and polyurethane urea tubes were made, respectively.

These tubes have a transparent appearance, and unlike the tubes of the above embodiments, the internal structure of the tube wall does not consist of a large number of continuous vacuole walls that make up the vacuole.
On the contrary, there were many parts with detailed structures. The tube was stiff, had poor elasticity, and lacked compliance.

For reference, the tube of Reference Example 3 was replaced with the tube of Example 2 described above.
．． The tube was transplanted into an adult mongrel dog in the same way as in case 5, but it was difficult to suture because it was hard and lacked elasticity, and blood leaked from the needle hole after suturing, and the transplant tube was occluded within a week.

11±-'' Stainless steel rods with outer diameters of 4 mm and 5+ sm, respectively, were added to the polyurethane and polyurethane urea solutions used in Examples 3 and 6 using a conventionally known method, that is, dip drying was repeated 10 times or more. and polyurethaneurea tubes were made.

These tubes have a transparent appearance, and unlike the tubes of the above embodiments, the internal structure of the tube wall does not consist of a large number of continuous vacuole walls that make up the vacuole.
On the contrary, there were many parts with detailed structures. The tube was stiff, had poor elasticity, and lacked compliance.

For reference, the tube of Reference Example 4 was replaced with the tube of Example 4 described above.
．． The tube was transplanted into an adult mongrel dog in the same manner as in case 6, but it was difficult to suture because it was hard and lacked elasticity, and blood leaked from the needle hole after suturing, and the transplant tube was occluded within a week.

1 unit U The solution used in Example 6 was extruded from the same annular nozzle as used in Reference Example 2, and water was similarly injected into the inside.

Due to the high viscosity of the solution, the back pressure increased to 8 Kg/-, and the outer surface of the obtained tube showed ejection spots, and the inner diameter was cyclic within the range of 3.0 to 3.5 mm. A change was observed.

The same solution as in Example 3 was extruded using the same nozzle and stainless steel rod, and after passing through a 200m+w dry section,
led to a coagulation bath.

The outer surface of the tube obtained in this example was smoother than that obtained in Example 3, and the outer diameter was determined by crossing the 1001 length tube in 5+sm increments. In this example, the 95% confidence interval is 7.55 to 7.6.
In contrast to 7.43 to 7.7 for those without a dry section.
75, and there was a large variation.

Fruit 1st 1-J The same solution as in Example 3 was extruded using the same nozzle and stainless steel rod, passed through a 100+*m dry section at a speed of 50 mm per second, and then introduced into a coagulation bath.

After coagulating in a running water bath overnight, the core rod was removed, boiled in boiling water for 30 minutes, and then dried at 45°C. The obtained artificial blood vessel was straight and had no bends, and there were no wrinkles or other deformations on the surface.

1 work 1 In Example 8, the speed of passing through the dry part was set to 32 per second.
A dried artificial blood vessel was obtained under the same conditions except that the diameter was 0 mm. It shrunk in the length direction, and many fine wrinkles were observed on the outer surface.

The polyurethane used in Example 3 was dissolved in dimethylacetamide to make a 10% solution. This solution was molded using a stainless steel rod previously covered with polyester mesh having five outer diameter layers and a 7-inch nozzle.

The polyester mesh was completely wetted with the polyurethane solution and embedded within the cross section, with no exposed surfaces on either the inside or outside.

The compliance value of this tube is 4% and 50
No irreversible morphological changes were observed even when internal pressure was applied at 0++vHg for one month.

A sample extruded under the same conditions as in Sub-Example 7 was introduced into a saturated saline solution, soaked overnight, washed with running water, and dried. The outer surface of the obtained tube had significantly fewer irregularities than that obtained in Example 7, and almost no undulations were observed even in an image magnified 5000 times.

[Effects of the Invention] According to the present invention, a medical tube having excellent long-term patency can be provided.

Claims (6)

[Claims] (1) A single-layer medical tube made of a polymer compound, characterized in that the internal structure of the tube wall constituting the tube is porous, and there is no skin layer on the inner surface of the tube wall. medical tube. (2) The internal structure of the tube wall is composed of a large number of continuously connected vacuole walls constituting the vacuole, and the inner surface of the tube wall is composed of the continuous and connected vacuole walls. A medical tube according to claim 1. (3) The medical tube according to claim 1 or 2, wherein the polymer compound is polyurethane and/or polyurethane urea. (4) By extruding a rigid core rod with a circular cross section from a circular orifice, a solution of a polymer compound is extruded through a gap slit between the orifice and the core rod so as to be spread over the entire circumferential surface of the core rod. . A method for manufacturing a medical tube, which comprises introducing the core rod into a coagulation bath to coagulate the polymer compound around the core rod, and then taking out the core rod. (5) The manufacturing method according to claim 4, wherein the polymer compound is polyurethane and/or polyurethane urea. (6) The manufacturing method according to claim 4 or 5, wherein the core rod is passed through a dry section immediately before being introduced into the coagulation bath.