WO1991019520A1

WO1991019520A1 - Microporous tubular prostheses

Info

Publication number: WO1991019520A1
Application number: PCT/NL1991/000105
Authority: WO
Inventors: Jan Feijen; Wouter Leonardus Joseph Hinrichs
Original assignee: Stichting Voor De Technische Wetenschappen
Priority date: 1990-06-21
Filing date: 1991-06-21
Publication date: 1991-12-26
Also published as: NL9001428A

Abstract

The invention relates to microporous tubular structures the wall of which is composed of a mixture of, on the one hand, a matrix of a (co)polymer having elastomer properties at body temperature and, on the other hand, a (co)polymer which is incorporated in said matrix and which is crystalline at body temperature or has a glass transition temperature of at least 37 °C with the proviso that mixtures of, on the one hand, polyester-urethanes or polyether-urethanes and, on the other hand, poly (lactic acid) are excluded. These tubular structures may be used as blood vessel prostheses.

Description

Microporous tubular prostheses

The invention relates to microporous tubular structures which can be used, for example, as blood vessel prostheses. Prostheses of this type are disclosed in Netherlands Patent

Application 82.02893. More particularly, this Netherlands patent application relates to prostheses based on, on the one hand, poly(L-lactic acid) and/or poly(D,L-lactic acid) and, on the other hand, segmented polyester-urethanes or polyether-urethanes, and specifically 5~95# by weight of poly(L-lactic acid) or poly(D,L-lactic acid) and 95-5# by weight of polyester-urethanes or polyether-urethanes. The tubular prostheses are produced with the aid of a stainless steel rod coated with polytetrafluoroethene, which rod is immersed in a polymer solution containing sodium citrate and then dried. This coating treatment for the rod is repeated until a tubular prosthesis having a desired wall thickness is obtained around the rod. A prosthesis obtained in such a way is then extracted for 5^_10 hours with distilled water and ethanol to remove the sodium citrate present in the prosthesis in order to obtain a certain degree of porosity and pore size respectively.

It has been found that for an optimum use the mechanical properties of the synthetic materials for tubular prostheses, such as blood vessel prostheses, must approach the mechanical properties of the specific tissue as closely as possible. After introduction into the body, the prosthesis must, by means of controlled fragmentation, make room for the tissue formed because of the ingrowth and overgrowth of endothelial cells and smooth muscle cells on the inside of the prosthesis and ingrowth of perivascular tissue on the outside of the prosthesis. During this process, the rate of fragmentation of the synthetic material must be matched to the regeneration of the replaced tissue in such a way that the mechanical properties of the implant and tissue combination are maintained as far as possible.

Surprisingly, it has been found that the rate of fragmentation of microporous prostheses produced from a (co)polymer which has elastomer properties at the body temperature can be increased by the addition of a (co) olymer which is incompatible with the elastomer and is crystalline at body temperature or has a glass transition temperature of at least 37°C. By this measure it is possible to increase the rate of fragmentation of, for example, a blood vessel prosthesis, produced from an, for example, very slowly degrading elastomer in such a way, that this rate is matched to the rate of regeneration of the blood vessel wall.

The fragments of the prosthesis material formed as a result of fragmentation must not cause any undesired tissue reactions. Preferably, according to the invention, prostheses are used which are built up from a mixture of, on the one hand, elastomer and, on the other hand, (co)polymers which are crystalline or have a glass transition temperature of at least 37°C, which fragment to form compounds which are not foreign to the body.

With regard to Netherlands Patent Application 82.02893. discussed above, the invention therefore provides, in addition to the great diversity of the elastomer and non-elastomer (co)polymers which can be used, the possibility that the polyester-urethane and polyether-urethane disclosed in said Netherlands patent application can be replaced by an elastomer material which fragments to form compounds which do not cause any undesired tissue reactions. The degradation products of polyether-urethane and polyester-urethane are possibly toxic and would be able to cause adverse effects in the long term.

Therefore, polymers which after degradation in the body do not form any products which cause adverse effects must preferably be used for both the elastic and the non-elastic component. These polymers are found in the groups: polyesters, polyorthoesters, polyanhydrides and poly(α-amino acids) and copolymers such as poly- depsipeptides. In this context in particular polymers are considered which are built up of units which also occur in the body. Examples of elastomers which can be used according to the invention and which degrade to form compounds which are not foreign to the body are elastomer (co)polymers built up from compounds which are not foreign to the body, such as L-lactic acid, D,L-lactic acid, glycolic acid, α-hydroxyvaleric acid, e-caprolactone and α-amino acids such as glycine and L-alanine. Non-elastic components which can be used are, inter alia, the homopolymers of the compounds mentioned. More particularly, the invention therefore relates to prostheses, for example blood vessel prostheses, which are elastic and microporous and are also composed of a mixture of two

(co)polymers which are preferably biodegradable to form compounds which are not foreign to the body, the one (co)polymer being elastic at body temperature, i.e. having elastomer properties, and the other

(co)polymer being crystalline at body temperature or having a glass transition temperature of at least 37°C, the latter (co)polymer acting as the controlling factor for the fragmentation of the pros- thesis. Specifically, it has been found that a prosthesis of this type fragments as a result of the mechanical stress exerted on the prosthesis. Simultaneously with the disintegration of the prosthesis, regeneration of, for example, a blood vessel takes place as a result of ingrowth and overgrowth, so that finally the prosthesis is taken up by a new blood vessel built up from tissue.

As indicated above, the initial degeneration of a prosthesis used, for example, as a synthetic blood vessel takes place via a fragmentation on the basis of the mechanical stress to which it is subjected and which in this example is caused by the arterial pulses. Because of the incompatibility of the two components, the structure of the prosthesis consists of an elastic matrix in which non-elastic domains are located. Under extreme stress, the non- elastic domains cause stress concentrations, as a result of which fracture will occur in the matrix. The weight ratio of the elastomer (co)polymer to the non- elastomer (co)polymer, from which the prostheses according to the invention are built up, may vary from 60:40 to 9 :1. preferably from 80:20 to 95:5. depending on, inter alia, the desired prosthesis characteristics and the materials used. With regard to the pore size and porosity of the prosthesis material according to the invention, it is pointed out that these latter parameters may vary from 5"200 μm and 10-95$. an preferably from 30~100 V™ an 70-90$. Moreover, the pore size itself may vary over the cross-section of the prosthesis wall. For example, the pore size on the lumen side is usually smaller (5~50 μm) than on the other side (outside) of the prosthesis (50-200 μm) , where direct ingrowth of body tissue must take place. The fragmentation mechanism is illustrated with reference to the experiment described below. The polyether-urethane/poly(L-lactic acid) combination specified in Netherlands Patent Application 82.02893 is used as reference. In this experiment use is made of two types of prostheses. The first type is built up from, on the one hand, a well-defined elastomer polyether-urethane (PEU) and, on the other hand, a (co)polymer which is crystalline at body temperature or has a Tg of at least 37"C, such as the polymers poly(e-caprolactone) (PCL) , polyglycine (PGLY) , poly(L-lactic acid) (Mw = 100,000 g/mol, (PLLA (MM)), poly(L-lactic acid) (Mw = 380,000 g/mol, (PLLA (HM)) and polysulphone (PSF) . The fragmentation mechanism is also illustrated with reference to prostheses which are built up from, on the one hand, the copolyester-ether elastomer Hytrel 4056 (Hytrel) and, on the other hand, a (co)polymer which is crystalline at body temperature or has a Tg of at least 37°C, such as the polymers poly(L-lactic acid) (Mw = 50,000 g/mol (PLLA (LM)), PLLA (HM) and PSF. It is emphasised that PSF is not biodegradable and is used solely to demonstrate that a fragmentation of the prosthesis nevertheless takes place. For comparison purposes the combinations of the elastomer polyester- urethane (PESU)/PEU and Hytrel/PEU and prostheses consisting of PEU alone and of Hytrel alone are also used, which, as can be seen from the experiment carried out, do not fragment.

The prostheses produced by means of dip coating are used as blood vessel prostheses to replace the ventral aorta in rats.

After an implantation period of 6 weeks, the prostheses are removed from the rat and treated with trypsin/collagenase in order to remove the ingrown tissue. The degree of fragmentation of the prostheses tested in this way can then be determined. Differential scanning calori etry (DSC) , dynamic torsion measurements and scanning electron microscopy (SEM) studies on a number of the blends used were carried out in order to demonstrate the incompatibility of the different polymers in the blend. Examples

Preparation of the blood vessel prostheses

The PEU/X prostheses are produced by means of dip coating.

This method is based on the application of a number of layers on a glass rod by dipping said rod alternately into a polymer solution and a non-solvent. For this operation the following procedure is followed:

A glass rod of the desired diameter is dipped for a few seconds into a polymer solution containing salt particles. The glass rod is then placed in a non-solvent, as a result of which the polymer coagulates. During this coagulation the salt particles are surrounded by a polymer matrix. When the polymer has completely coagulated, the prosthesis is dried, first using tissue paper and then in the air, until the outside no longer feels damp. The glass rod is turned round and the following layer is applied in the same way. This procedure is repeated until the desired thickness of the prosthesis wall has been achieved. By then removing the salt from the prosthesis by means of extraction, a prosthesis is obtained which has a porous structure. The Hytrel/X prostheses were produced by means of a comparable procedure. In this case also, various porous layers were applied to a glass rod by dipping said rod several times into a polymer solution containing salt particles. However, the method of precipitation was different. In the case of the production of the Hytrel/X prostheses, the glass rod dipped into the polymer solution containing salt particles was exposed to the air. Precipitation of the polymer took place as a result of evaporation of the solvent.

The pore size and the porosity of the prosthesis can be varied by changing the particle size of the salt particles and/or the salt and polymer concentrations of the dip solutions. A porosity distribution and pore size distribution can be obtained by using different solutions for the successive dip steps.

In the experiment according to the invention, prostheses are produced having the composition PEU/X = 9/1 (w/w) , where X = PEU, PESU, PCL, PGLY, PLLA (MM), PLLA (HM) and PSF, and Hytrel/X = 9/1 (w/w) where X = PEU, Hytrel, PLLA (LM) , PLLA (HM) and PSF. The abovementioned abbreviations relate to: Hytrel = copolyester-ether, the commercial product Hytrel 4056; PEU = polyether-urethane, the commercial product Estaan 5714F1; PESU = polyester-urethane, the commercial product Estaan

5707F1; PCL = poly(e-caprolactone) (molecular weight: 70,000); PGLY = polyglycine (molecular weight: 2,000); PLLA (LM) = poly(L-lactic acid) (molecular weight: 50,000); PLLA (MM) = poly(L-lactic acid) (molecular weight: 100,000); PLLA (HM) = poly(L-lactic acid) (molecular weight: 380,000); PSF = polysulphone (Union Carbide P3500 having the empirical formula (C₆H_₁(CH₃)₂C₆H₄0C₆H_itS0₂C₆H_/(0)_n) . A porosity distribution and pore size distribution is introduced by using two dip solutions. The compositions of these solutions are given in Table A below. The diameter of the glass rod was 1.3 mm or 1.5 mm. The temperature of the dip baths was controlled such that a uniform thin layer could be applied by the dip procedure. In the case of the production of the PEU/X prosthesis, the temperature varied per composition and was between 40 and 60°C. In the case of the production of the Hytrel/X prostheses, said temperature was always about 20^βC. In the case of the production df the PEU/X prostheses, coagulation took place during the first 10 seconds in an ethanol/water mixture having a ratio of 10/0, 9/1 or 8/2 (v/v) at room temperature and the glass rod was then transferred to ethanol/water = 7/5 (v/v) , likewise at room temperature (for a minimum of 5 minutes) . In the case of the production of the Hytrel/X prostheses, the glass rod was exposed to the air for 1 to 3 minutes. After the desired number of layers had been applied, the PEU/X prostheses were placed in ethanol/water = 7/5 (v/v) for 30 minutes and the Hytrel/X prostheses were exposed to the air for 20 minutes. The salt was then extracted from the prosthesis using distilled water at 4θ-50°C over a period of 2 hours. The prosthesis was then rinsed twice with distilled water, removed from the glass rod and stored in ethanol. TABLE A

PEU/X prostheses: X: PaEU PaESU/ PLLA (MM)^f PLLA (HM)^f

PCL/PGLYV PSF

Lumen side:

Poly. cone.^a 7.5-8.0 7-1 7.0 Solvent 7/1/0 or 1/0/0 10/2/5 10/2/5 Salt conc.^c 50 33 11.5 Particle size salt < 38 < 38 < 38 Number of layers 2 to 3 3 3

Outside:

Polym. cone.^a 6.6 6.6 6.6 Solvent^b 5/1/0 or 5/0/08/1/3 8/1/3 Salt conc.^c 16 Particle size salt^d 63-106 Number of layers

2

Hvtrel/X prostheses: X: Hytrel^g/PLLA (LM)/ PEU PLLA (HM)/PSF

Lumen side:

Polym. cone.^a 7 7 Solvent" 0/0/1 0/1/4 Salt conc.^c 85 85 Particle size salt^d < 38 < 38 Number of layers 2 to 3 2

Outside:

Poly . cone.^a 4.5 4.5 Solvent¹¹ 0/0/1 0/1/4 Salt conc.^c 81 81 Particle size salt^d 63-106 63-106 Number of layers 5 to 8 6 a: Polymer concentration in g/100 ml b: Ratio of N,N-dimethylformamide (DMF)/tetrahydrofuran

(THF)/CHC1₃ (v/v/v) c: Concentration of sodium citrate in g/100 ml d: Particle size in μm e: PGLY does not dissolve in DMF/THF mixtures. The PGLY is dispersed (d < 5 μm) over the PEU matrix f: PLLA (MM) and PLLA (HM) are sparingly soluble in DMF/THF mixtures but readily soluble in CHC1₃. Just before the dip procedure, a solution of PLLA (MM) or PLLA (HM) in CHC1₃ is added dropwise to a solution of PEU in DMF/THF, while stirring very well, until a solution having the desired composition is obtained, g: Prostheses consisting of PEU alone or of Hytrel alone are indicated by PEU/X, X = PEU and, respectively, Hytrel/X,

X = Hytrel.

Characterisation of prostheses

The prostheses obtained using the dip method described above were checked for porosity, cracks, uniform wall thickness, pin holes and layer adhesion using light microscopy. The wall thickness was also determined using light microscopy.

The porosity was calculated on the basis of the weight and the volume of the prostheses and the pore size distribution was determined using scanning electron microscopy.

The elasticity of the prostheses in the longitudinal direction was determined as follows. Shortly before the measurement, the prosthesis removed from ethanol was rinsed with distilled water and fixed between the clamps of an Instron tensile tester. The prosthesis, having a length of 1 cm, was stretched 100$ using a pulling speed of 0.5 cm/min. The force measured during this stretching treatment gave the F-100$ value.

In a number of cases tensile/stretching measurements were also carried out in the radial direction. In this case, the prosthesis having a length of 0.5 cm was stretched at a speed of

0.5 cm/min. The force measured for stretching by 0.5 cm gave the

FR(0.5) value.

With regard to the component X in the PEU/X prostheses and the Hytrel/X prostheses, the data given below are provided for illustration, Tm being the melting point, Tg the glass transition temperature, the term "n.dg" denoting poorly degradable, the term

"dg" denoting degradable, the term "el" denoting elastic and the term "n.el" denoting non-elastic .

With regard to the prostheses obtained using the dip method described above, the following data are provided in Table C below.

TABLE C- Composition

PEU

PEU/PESU = 9/1 PEU/PCL = 9/1 PEU/PGLY = 9/1 PEU/PLLA (MM) = 9/1 PEU/PLLA (HM) = 9/1 PEU/PSF = 9/1 Hytrel

Hytrel/PLLA (LM) = 9/1 Hytrel/PLLA (HM) = 9/1 Hytrel/PSF = 9/1 Hytrel/PEU = 9/1

Compatibility

Various techniques were used on a number of blends to obtain an indication with respect to the degree of mixing at a molecular level of PEU and X or of Hytrel and X in these blends. Both porous prostheses and dense (non-porous) films were used. The dense films were produced by means of evaporation of polymer solutions.

The crystallinity of the component X in the PEU/X prostheses where X = PCL, PLLA (MM) and PLLA (HM) and in the Hytrel/X prostheses where X - PLLA (LM) and PLLA (HM) relative to pure X was determined using a DuPont 990 thermal analyser (DSC) using a rise in temperature of 10°C/min. The thermograms of the PEU/PLLA (MM) and the PEU/PLLA (HM) prostheses showed two melting peaks between l6θ°C and l8θ°C. Apparently two types of crystal lattices of the PLLA having different melting points are formed in the PEU matrix. The DSC measurements also show that the crystallinity of the component X in the various blends has decreased only slightly compared with the pure component. This indicates a high incompatibility of the polymers in the mixture.

TABLE D

PCL (100$ crystalline) PCL (starting material) PEU/PCL prosthesis

PLLA (100$ crystalline) PLLA (MM) (starting material) PEU/PLLA (MM) prosthesis

PLLA (100$ crystalline) PLLA (HM) (starting material) PEU/PLLA (HM) prosthesis

Hytrel/PLLA (HM) prosthesis

PLLA (100$ crystalline) PLLA (LM) (starting material) PEU/PLLA (LM) prosthesis

a: theoretical value

Dynamic torsion measurements on the prostheses and dense films (20 x 6 x 0.5 mm) having the composition Hytrel/X = 9/1 where X = Hytrel, PLLA (LM) , PLLA (HM) , PSF and PEU were carried out using a Myrenne torsion balance. The measurements were carried out using a torsion frequency of 1 Hz and a rise in temperature of 1 C/min. A glass transition temperature of -60°C was measured for all materials tested. It can be concluded from this that the amorphous fractions of the polymers have not mixed or have virtually not mixed in the various blends.

By placing the Hytrel/X blends in a selective solvent it was possible to remove the X additive from the blend and to render the phase separation visible with the aid of SEM. Dense films were used in these experiments because with porous structures it was not possible to differentiate between the pores which were already present prior to the solvent treatment and the pores which formed as a consequence of the selective removal of the additive. Dioxane was found to be a suitably selective solvent for the removal of the various X additives from the Hytrel/X blends. Cross-sections of the dense films having the composition Hytrel/X = 9/1. where X = Hytrel, PEU, PLLA (LM) , PLLA (HM) and PSF, obtained using microtomy at -120^βC, were incubated for 15~30 minutes in dioxane. The samples were then dried and studied with the aid of SEM. It was observed that the cross-sections of the dioxane-treated samples having the composition Hytrel/X = 9/1 where X = PEU, PLLA (LM) , PLLA (HM) and PSF were porous and the cross-section of the dioxane-treated sample consisting of Hytrel alone was not porous. For the sake of completeness, it is pointed out that the cross-sections of samples not treated with dioxane showed no pores. These measurements demonstrate that the polymers in the blends used have not mixed with one another on the molecular level or have done so only to a very slight extent.

In vivo evaluation The abovementioned prostheses having a length of 1 cm were used for a period of 6 weeks as blood vessel prostheses to replace the ventral aorta of a rat (N = 3) • After explantation, the PEU/PLLA (MM) prosthesis and one of the three Hytrel/PLLA (LM) prostheses were found to be highly dilated, which is an indication of advanced fragmentation of the synthetic material. The other prostheses used had a normal appearance after explantation. The ingrown tissue was removed with the aid of a combined trypsin/collagenase treatment. More particularly, the prostheses obtained after explantation were stored in ethanol and rinsed thoroughly with distilled water before the trypsin/collagenase treatment. 5 ml of 10$ trypsin/EDTA were then added to the prostheses rinsed in this way and the prostheses were treated for 24 hours at 37°C in a shaking bath. After extraction with distilled water, the prostheses were finally treated for 2 to 4 days with 5 ml of 0.3$ collagenase in PBS in a shaking bath at 37°C.

After the collagenase/trypsin treatment the PEU/X prostheses where X = PSF, PCL and PLLA (MM) were found to be highly fragmented: the prostheses showed large cracks and crumbled under slight tension. The PEU/PLLA (HM) and the PEU/PGLY prostheses were less highly fragmented: the prostheses appeared macroscopically intact, but the tensile strength was reduced to very low values. The PEU/X prostheses where X = PEU and PESU were hardly fragmented: a slight weakening of the prostheses could be determined with the aid of tensile/stretching measurements (see Table E) . The Hytrel/PLLA (LM) prostheses were substantially fragmented: the dilated prosthesis had fallen apart into three pieces and in the case of the other two prostheses large cracks were to be seen and the tensile strength was reduced to very low values. The Hytrel/PLLA (HM) and the Hytrel/PSF prostheses appeared unchanged. However, it was possible to detect fragmentation by means of measurements using the tensile tester (see Table E) . In the case of the Hytrel/Hytrel and the Hytrel/PEU prostheses, no fragmentation could be found. The mechanical properties of the prostheses before and after implantation were virtually identical. For the sake of completeness it is pointed out that a combined trypsin/collagenase treatment had no effect on the mechanical properties of the non-implanted prostheses. TABLE E

PEU/X prostheses: expl.^a tri.

PEU +

PESU +

PCL +

PGLY +

PLLA (MM)

PLLA (HM)

PSF

Hytrel/X prostheses:

X expl . ^a tri . -coll . ^b FR(0.5)_b/ TS_b/TS_s S_b/S_ε

FR(0.5) ,

Hytrel 1 .90/1.80 2.10/1.80 6.5/5 PLLA (LM) na/nd na/nd na/nd PLLA (HM) na/na 2.10/0.75 3.6/1. PSF 4.00/na 4.70/2.30 7.5/3 PEU 5.80/5.10 10.5/9.0 8.5/8.0

nd not determined a) unchanged dilated b) not or hardly fragmented decrease in the mechanical properties tensile strength was reduced to very low values

: very extensively fragmented; the prostheses show cracks c) FR(0.5)_b/FR(0.5)_a: FR(0.5) value in Newton before implantation/FR(0.5) value in Newton after implantation d) TS TS, maximum tensile strength in Newton before implantation/maximum tensile strength in Newton after implantation e) S./S, stretch at maximum tensile strength in mm before implantation/stretch at maximum tensile strength in mm after implantation

Claims

CLAIMS 1. Microporous tubular structures, characterised in that the wall is composed of a mixture of, on the one hand, a matrix of a (co)polymer having elastomer properties at body temperature and, on the other hand, a (co)polymer which is incorporated in said matrix and which is crystalline at body temperature or has a glass transition temperature of at least 37°C, with the proviso that mixtures of, on the one hand, polyester-urethanes or polyether- urethanes and, on the other hand, poly(lactic acid) are excluded.

2. Microporous tubular structures according to Claim 1, characterised in that the (co)polymer which is incorporated in the elastomer matrix and which is crystalline at body temperature or has a glass transition temperature of at least 37°C is a biodegradable (co)polymer.

3« Microporous tubular structures according to Claim 1 or 2, characterised in that the (co)polymer which is incorporated in the elastomer matrix and which is crystalline at body temperature or has a glass transition temperature of at least 37"C is a biodegradable (co)polymer which is made up of units which occur in the body and of which the degradation products do not give rise to any adverse reactions.

4. Microporous tubular structures according to Claim 1, 2 or 3, characterised in that the (co)polymer which is incorporated in the elastomer matrix and which is crystalline at body temperature or has a glass transition temperature of at least 37°C is a biodegradable (co)polymer selected from the group consisting of polyesters, polyorthoesters, polyanhydrides and poly(α-amino acids) and copolymers such as polydepsipeptides.

5. Microporous tubular structures according to Claim 1, 2, 3 or 4, characterised in that the (co)polymer which is incorporated in the elastomer matrix and which is crystalline at body temperature or has a glass transition temperature of at least 37°C is a biodegradable (co)polymer chosen from the group consisting of (co)polymers built up from L-lactic acid, D,L-lactic acid, glycolic acid, α-hydroxyvaleric acid, €-caprolactone and α-amino acids such as glycine and L-alanine.

6. Microporous tubular structures according to Claim 1, characterised in that both the (co)polymer which has elastomer properties at body temperature and the (co)polymer which is crystalline at body temperature or has a glass transition temperature of at least 37°C are biodegradable (co)polymers.

7. Microporous tubular structures according to Claim 6, characterised in that the (co)polymer which is incorporated in the elastomer matrix and which is crystalline at body temperature or has a glass transition temperature of at least 37°C is a biodegradable (co)polymer chosen from the group consisting of polyesters, polyorthoesters, polyanhydrides and poly(α-amino acids) and copolymers such as polydepsipeptides.

8. Microporous tubular structures according to Claim 6 or 7. characterised in that the (co)polymer which is incorporated in the elastomer matrix and which is crystalline at body temperature or has a glass transition temperature of at least 37°C is a biodegradable (co)polymer which is built up from units which occur in the body and of which the degradation products give- rise to no adverse reactions.

9. Microporous tubular structures according to Claim 6, 7 or 8, characterised in that the (co)polymer which is incorporated in the elastomer matrix and which is crystalline at body temperature or has a glass transition temperature of at least 37°C is a biodegradable (co)polymer chosen from the group consisting of (co)polymers built up from L-lactic acid, D,L-lactic acid, glycolic acid, α-hydroxyvaleric acid, €-caprolactone and α-amino acids such as glycine and L-alanine.

10. Microporous tubular structures according to Claim 1, 6, 7. 8 or 9. characterised in that the elastomer (co)polymer is built up from units which occur in the body and of which the degradation products give rise to no adverse reactions.

11. Microporous tubular structures according to Claim 1, 6, 7. 8 or 10, characterised in that the elastomer (co)polymer is chosen from the group consisting of polyesters, polyorthoesters, polyanhydrides and poly(α-amino acids) and copolymers such as polydepsipeptides.

12. Microporous tubular structures according to Claim 1, 6, 7. 8, 9. 0 or 11, characterised in that the elastomer (co)polymer is chosen from the group consisting of (co)polymers built up from L-lactic acid, D,L-lactic acid, glycolic acid, α-hydroxyvaleric acid, e-caprolactone and α-amino acids such as glycine and L-alanine.

13« Microporous tubular structures according to one or more of Claims 1-12, characterised in that said structures are used as blood vessel prostheses.

14. Microporous tubular structures according to one or more of Claims 1-13. characterised in that the weight ratio of the elastomer (co)polymer to the non-elastomer (co)polymer is 60:40 to 99:1-

15. Microporous tubular structures according to Claim 14, characterised in that the weight ratio of the elastomer (co)polymer to the non-elastomer (co)polymer is 80:20 to 95:5-

16. Microporous tubular structures according to one or more of Claims 1-15. characterised in that the pore size is 5-200 urn and the porosity is 10-95$.

17. Microporous tubular structures according to Claim 16, characterised in that the pore size is 30-100 μm and the porosity is 70-90$.