Biodegradable stent with adjustable degradation rate

BACKGROUND OF THE INVENTION

Field of the Invention

The present invention relates to a method of treating blood vessel diseases with bioresorbable medical devices made of biodegradable polymers, in particular stents.

Description of the state of the art

The present invention relates to radially expandable endoprostheses adapted for implantation in the light of the anatomical structures of the body. The term "endoprosthesis" means a synthetic medical device placed inside the body. The term "light" means a cavity in the interior of a coil-shaped (tube) organ, such as a blood vessel. An example of such an endoprosthesis is a stent. In general, stents are cylinder-shaped devices whose function is to keep the light open, and sometimes to expand a section of the blood vessel or other hollow anatomical structures such as the urinary tract and bile ducts. Stents are often used to treat narrowing caused by atherosclerosis. The term "stenosis" means the narrowing or contraction of the tube or opening of the anatomical structure in the body. This method of treatment involves strengthening the wall of the blood vessel with a stent, which prevents restenosis after angioplasty in the vascular system. The term "restenosis" means the recurrence of narrowing of the blood vessel or heart valve after successful treatment (e.g., balloon angioplasty, stenting or valvuloplasty).

Stents usually consist of scaffolding in the form of a mesh interconnected structural elements, i.e. struts, made of wire, tubes or sheets of material rolled into a cylinder shape. This "scaffolding" bears this name because it keeps the light open, and if necessary stretches the walls of the duct. Usually, stents can be compressed (squeezed) or folded on the catheter in such a way as to allow their insertion and expansion at the site requiring treatment.

The introduction consists in placing the stent with a catheter in the lumen of a small duct and delivering it to the place requiring treatment. Stretching involves developing the stent to a larger size after placing it in the desired place. Mechanical intervention using stents has reduced the incidence of restenosis compared to balloon angioplasty. Still, restenosis is still a major problem. In the case of restenosis in the episode with a stent on, treatment can be difficult due to the limited number of options compared to lesions treated only with the balloon technique.

Stents are used not only as a kind of mechanical intervention, but also as carriers of drugs in biological therapies. Biological therapy involves the use of drug-releasing stents to deliver the drug substance to its destination. The drug substance may also alleviate the unwanted biological response to the presence of the stent in the body. Achieving an effective concentration at the site requiring treatment requires systemic drug administration, which is often associated with adverse and even toxic effects. Delivering the drug to the destination is the preferred method of treatment because of the possibility of giving lower total amounts than for systemically administered drugs, while at the same time obtaining a higher concentration at a specific site. Therefore, delivering the drug to the destination causes fewer side effects and allows to achieve better results.

A drug releasing stent can be made by covering the surface of a scaffold made of metal or polymers with a polymer carrier containing an active or bioactive drug or agent. Also, a polymer scaffold can serve as a carrier for an active drug or agent.

Such a stent must meet many of the required mechanical parameters. It must have adequate radial strength sufficient to overcome structural loads, namely radial compressive forces acting on the stent supporting the vessel walls. After expansion, the stent must ensure adequate patency of the vessel lumen within the required time, despite the various forces exerting pressure on it, including cyclical loads caused by beating hearts. In addition, the stent must be flexible enough and to some extent resistant to breaking.

The use of stents made of biostable or non-erosive materials, such as metals, has gained the status of "standard in clinical practice" for such procedures as percutaneous coronary intervention (PCI), as well as for peripheral applications (e.g. in the superficial femoral artery), because the ability to prevent early and late reposition and to reduce the diameter of the stent after implantation has been demonstrated.

To cure a diseased blood vessel, the presence of a stent is only required for a limited time. The permanent presence of the implant in the vessel has some disadvantages, such as the biological incompatibility of the stent and vessel, and the risk of embolic incidents. They can be minimized by making stents from materials that are eroded or degraded as a result of exposure to internal conditions. In this way, the eroded part of the stent may disappear from the implantation site after treatment, leaving the vessel repaired. Stents produced from biodegradable, bioresorbable and / or bioerodible materials, such as bioresorbable polymers, can be constructed to be completely eroded only after the required period of use.

Like a permanent stent, the biodegradable stent must meet the required mechanical parameters depending on the passage of time. For example, it must ensure patency for the minimum necessary time. It is also important for the biodegradable stent to completely disintegrate at the implantation site within a certain period of time. Biodegradable material that meets the required mechanical parameters may not degrade in the necessary or desired time. In addition, the necessary or desired degradation time varies depending on where the stent is used, i.e. in the coronary vessels or perimeter.

EP 1 184 008 A1 describes: a stent intended for use in the treatment of a diseased section of a blood vessel consisting of a polymer scaffold made of poly (L-lactide) (PLLA), the degree of crystallinity of poly (L-lactide) being varied in the range of 15 to 60%, and the average molecular weight of poly (L-lactide) is 60,000 to 300,000, depending on the polystyrene standard. US 2007/253999 and WO 2007/146354 A2 disclose biodegradable stents made of poly (L-lactide).

BENDIX D: In the work "Chemical synthesis of polylactide and its copolymers for medical applications", POLYMER DEGRADATION AND STABILITY, BARKING, GB, it was shown that Llactide monomers contained in PLLA are considered impurities causing rapid (fast) decomposition during heat treatment.

SUMMARY OF THE INVENTION

Various embodiments of the present invention include a stent for use in the treatment of a diseased segment of a blood vessel consisting of a polymer scaffold made of poly (L-lactide) (PLLA) and L-lactide monomers at a mass percentage from 0.05 wt% to 0.5 wt % mixed, dispersed or dissolved in poly (L-lactide), with the degree of crystallinity of poly (L-lactide) ranging from 20 to 50%, and the average molecular weight of poly (L-lactide) is between 60,000 and 300,000, depending on the polystyrene standard.

Additional embodiments of the present invention include a method of treating a diseased section of a blood vessel comprising: expansion in the diseased section of the blood vessel of a bioresorbable polymer stent consisting of a body containing a scaffold made of expanding elements, the body being made of poly (L-lactide) and 0.05-0.5% L-lactide monomers, where said stent supports the wall of the blood vessel expansion diameter or close to it for a specified period of time, followed by a reduction in the radial strength of the stent and the stent stops supporting the vessel wall, and the scaffold separates and is completely absorbed by the human body.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts an example stent.

FIG. 2 is a diagram of a PLLA stent degradation process containing various concentrations of L-lactide monomer in vitro.

FIG. 3 is a diagram of degradation rate constants versus concentration of L-lactide monomer prepared based on the data presented in FIG. 2.

DETAILED DESCRIPTION OF THE INVENTION

In general, coronary arteries are arteries that depart from the aorta and supply the heart with oxygenated blood. Peripheral arteries generally mean vessels outside the heart and brain.

In both coronary artery disease (coronary artery disease) and peripheral artery disease, the arteries harden and narrow, which causes blood flow to be restricted. Coronary arteries reduce the blood supply to the heart, whereas peripheral arteries reduce the blood supply to the kidneys, stomach, arms, legs and feet. The narrowing is caused by the accumulation of cholesterol and other substances on the inner wall of the vessel, which are called atherosclerotic plaques. Such narrowing of the vessel sections is often called atherosclerosis. Arterial disease also relates to recurrence of stenosis after angioplasty treatment. Although there are probably several mechanisms underlying the changes leading to arterial restenosis, one of the most important is the development of an inflammatory reaction that stimulates tissue proliferation around the site of angioplasty. The inflammatory reaction can be caused by the expansion of the balloon used to open the lumen of the vessel, or by the presence of a stent, which is in itself a foreign body.

Embodiments of the present invention may be used in the treatment of coronary artery disease and peripheral vascular disease, including lesions of the superficial femoral artery, iliac artery and carotid artery. Embodiments may additionally be used for various types of stents, such as self-expanding stents and balloon-expandable stents. Embodiments can additionally be used for a variety of stent designs, including tubular, wire and braided mesh scaffolding.

In embodiments of the present invention, the stent consists of a variety of cylindrical rings connected or fastened by welding elements. After expansion in the vessel section, the cylindrical rings transfer loads and support the walls of the blood vessel, maintaining the diameter of the expansion or diameter depending on the cyclic forces acting inside the vessel. Load bearing means holding the weight exerted by inward radial forces. Structural elements, such as bonding or expansion elements, do not have load-bearing properties and their function is to maintain connections between the rings. For example, a stent may consist of a scaffolding in the form of a mesh of interconnected structural or strut elements.

FIG. 1 depicts the appearance of an exemplary stent 100. In some embodiments, the stent may consist of a body, skeleton, or scaffold in the form of a mesh of interconnected structural elements 105. The stent 100 may be formed of a tube (not shown). FIG. 1 presents features typical of many stent models, including cylindrical rings 107 connected by welding elements 110. As mentioned above, cylindrical rings transfer loads exerting a radially directed force to support the vessel walls. Basically, the task of the bonding elements is to connect cylindrical rings to each other.

The structural model shown in FIG. 1 is only an example and serves to illustrate the basic structure and features of the stent. Such a stent, like stent 100, can be made from a tube or polymer plate by rolling and welding the plate to form a tube. The tube or plate can be formed by extrusion or injection molding. A stent model, such as in FIG. 1, can be made of a tube or plate using techniques such as laser cutting or chemical etching. Then the stent can be rolled up on a balloon or catheter so that it can be inserted into the lumen of the vessel.

The dominant mechanism of degradation of biodegradable polymers is chemical hydrolysis of the skeleton, which is not resistant to hydrolysis. In the case of bulk eroding polymers, the polymer undergoes chemical degradation, and material loss occurs in the entire volume of the polymer. As the degradation proceeds, the molecular weight of the polymer decreases. After the molecular weight decreases, the mechanical properties deteriorate, followed by erosion, i.e. weight loss. Deterioration of mechanical properties ultimately causes a loss of mechanical integrity, which results in the device breaking up into pieces. As a result of the action of enzymes and the metabolism of the resulting fragments, there is a rapid loss of polymer mass.

The stent object of this invention used to treat arterial vascular disease after implantation has time-dependent properties that allow treatment and repair of a diseased segment of a vessel. In particular, molecular weight, mechanical properties, mechanical integrity and mass of stent change over time. After expansion in the diseased segment of the artery, the stent supports the vessel wall for some time, expanding its lumen to a larger diameter. As a result of the molecular weight reduction, radial strength is reduced to a level where the stent can no longer support the vessel wall in a given segment. The term "radial strength" of a stent is the pressure at which irreversible deformation of the stent occurs. After reducing the radial strength, mechanical integrity is gradually weakened.

Mechanical integrity refers to the size, shape and consistency of the structural elements of the stent. For example, shape generally means the tubular shape of a stent formed by cylindrical rings joined by structural bonding elements. The loss of mechanical integrity begins after the appearance or enlargement of cracks in the structural elements of the stent as a result of chemical degradation (reduction of molecular weight). Further loss of mechanical integrity occurs after breaking or loss of connections between structural elements.

Initially, the need for any stent arises from the need to provide mechanical support to maintain patency or to maintain the lumen of a vessel with or near an expansion diameter. The patency provided by the stent allows for a favorable reconstruction of the patented segment of the vessel at an increased diameter due to expansion. By maintaining patency of the patented segment of the vessel at this stage, the stent prevents its adverse reconstruction. Reconstruction generally means changes in the structure of the vessel wall resulting in a strengthening of its load resistance in such a way that the vessel wall in the patented section can maintain an increased diameter even without stent support. To allow permanent beneficial reconstruction of the vessel wall, it is necessary to maintain patency for some time.

During this time, the stent inhibits or prevents the natural pulsating movement of the vessel. The stent design prevents the diameter of the stent from being reduced after implantation and maintains the round shape of the lumen of the vessel, while the vessel undergoes remodeling and is modeled to fit the diameter of the patented segment (which means beneficial remodeling of the vessel wall). Premature reduction of the diameter of the stent after implantation, before the vessel is properly shaped, can cause its unfavorable reconstruction. This means that the vessel wall is formed in the patented section in such a way that its diameter becomes significantly smaller than the original diameter, e.g. 50% smaller than the original diameter after the stent has expanded.

As the polymer degradation process of the stent is made, the radial strength of the stent decreases, and the load on the vessel is gradually transferred from the stent to the remodeled vessel wall. Reconstruction of the vessel wall continues after the loss of radial strength of the stent. Before the stent loses its mechanical integrity, it is advisable to embed the structural elements of the stent into the vessel wall through the endothelial layer. The stent then breaks off, allowing the vessel to pulsate. The process of remodeling the vessel wall progresses after restoration of vessel motility as a result of pulsating movements. After all, the stent is completely eroded, leaving a repaired vessel with an increased diameter, which can pulsate in a similar or the same way as the correct section of the vessel.

Poly (L-lactide) (PLLA) is an attractive material due to its relatively high strength and rigidity at a human body temperature of about 37 ° C. Because the PLLA glass transition temperature is about 60-65 ° C (Medical Plastics and Biomaterials Magazine, March 1998), the polymer maintains stiffness and hardness at human body temperature. Due to this property, the stent maintains the ability to maintain the lumen of a vessel with or near an expansion diameter without significantly reducing the diameter of the stent after implantation.

The in vitro degradation time of PLLA is a maximum of 3 years (Medical Plastics and Biomaterials Magazine, March 1998; Medical Device Manufacturing & Technology 2005). This is the time to allow total weight loss of a polymeric structure, such as a stent. The in vivo degradation time is shorter and depends on the type of experimental animal model. In addition to the erosion profile, the PLLA stent has appropriate mechanical parameters (e.g. strength) and molecular weight. As shown above, the dependence of the mechanical properties and mechanical integrity of the stent on time is important in the treatment of diseased vessels. The PLLA degradation pattern does not generally coincide with the parameters required or desired for specific clinical applications. For example, the preferred degradation time for a stent made of PLLA is about two years (e.g., 22 to 26 months) in coronary vessels, about eighteen months (e.g. 16-20 months) in peripheral vessels (e.g. superficial femoral artery) and less than a year in nasal vessels. For coronary and peripheral arterial implantation, radial endurance should be maintained for one to six months, with radial endurance indicating that the stent maintains the lumen of the vessel in a patented segment at least 50% in diameter after expansion. In the case of coronary and peripheral artery implantation, the loss of mechanical integrity should not start earlier than after 2 to 4 months, with the beginning of the loss of mechanical integrity indicating the separation of structural elements.

It is advisable to adapt the stent from PLLA for various applications in such a way that the degradation process proceeds properly for each application, without a significant change in the composition of the stent.

Embodiments of the present invention relate to the modification of the time-dependent scheme of degradation of stent from PLLA as a result of the inclusion of L-lactide monomer (LLA) in the body of the stent made of

PLLA.

Embodiments of this stent may include a body or scaffold made of PLLA with the addition of a target small amount of LLA monomer. The body can be made of a target amount of LLA, and the rest of the body can be 100% PLLA. In addition, the body can be made of a target amount of LLA and additional components, and the rest of the body can be 100% PLLA. The body may contain PLLA in a mass percentage of at least 95 wt%, and the rest of the body may contain a target amount of LLA and optionally additional components. Additional components may be drugs, polymers or fillers such as bioceramic particles. The present invention includes embodiments excluding additional ingredients or at least one of the additional ingredients mentioned above.

In additional embodiments, the PLLA stent body may additionally or alternatively contain d-lactide, mesolactide, glycolide, lactic acid or oligomers thereof with a number average molecular weight (Mn) of less than 1000 g / mol. The concentration of these additional monomers may be the same as for L-lactide.

The stent may additionally contain a coating applied to the body or scaffolding. In one embodiment, the coating may consist of a mixture of polymer and drug. For example, the coating may be made of poly (DL-lactide) and the drug may be a substance with antiproliferative activity, such as everolimus. The coating may not contain LLA monomer, except in quantities resulting from accidental migration or diffusion of LLA into the coating.

Low molecular weight PLLA oligomers can also accelerate the rate of degradation and thus modify the degradation pattern. This is primarily due to the presence of acid end groups that act as catalysts to increase the rate of PLLA degradation. Thus, the larger the oligomer molecule, the greater the oligomer weight fraction in PLLA must be. Therefore, in order to obtain an oligomer-like effect, a much smaller weight proportion of LLA monomer is required than for a given oligomer. The high proportion of oligomer by weight can have a negative effect on the mechanical properties of the stent.

The exact and qualitative effect of LLA monomers in PLLA on the degradation pattern is unknown. For example, it is not known how much monomer is needed to achieve the desired degradation pattern. This can be at least partly attributed to the fact that the degradation scheme of a stent made of semi-crystalline degradable polyester, such as PLLA, is a complex derivative of several properties of the material and body of the stent. These properties include the specific rate of polymer hydrolysis (i.e. polymer chain skeleton disruption reaction), degree of crystallinity, morphology (size and distribution of crystal domains in the amorphous matrix), molecular weight (determined on the basis of natural viscosity and number average molecular weight) and parameters of the stent body (model, dimensions of the expansion elements).

In general, the strength and fracture toughness of the semi-crystalline polymer may not be sufficient to provide adequate and safe treatment of the blood vessel. The production process of the stent object of the present invention includes machining that increases the strength and fracture strength of the end product in the form of a stent. Thanks to this treatment, the material gains some features affecting the degradation scheme, e.g. crystallinity, morphology. Strength and fracture toughness are increased by obtaining a biaxial orientation of the polymer structure in the annular or circular system and in the axial direction, a certain degree of crystallinity and the presence of small dispersed crystallites (grains).

The stent is made of a PLLA extruded polymer tube that has been stretched radially and extended in the axial plane to achieve the desired orientation. The polymer tube is stretched by blow molding, radially expanded by 200% to 500%, and extended by 20% to 200% in the axial plane. Extruded PLLA tubes can be stretched in the axial plane from 100% to 400%. The stent is formed from a polymer tube in a stretched state by laser cutting.

In addition, blow molding treatment disperses small crystallites into an amorphous matrix. Before stretching, the tube is heated to a temperature of 65 ° C to 75 ° C to promote the formation of smaller crystallites that increase fracture resistance. After stretching, the tube is cooled below the glass transition temperature (Tg) to prevent further crystal growth. The degree of crystallinity is 20-50%. At less than 20% crystallinity, the stent body may not be strong enough, while at more than 50% crystallinity, the stent body may be too brittle. The average number molecular weight (Mn) of the material from which the scaffolding of the final product is made (expressed in g / mol) is from 60,000 to 300,000, or more precisely from 80,000 to 200,000.

An example of an expansion element may have a rectangular cross-section, e.g. 140 x 140 μm to 160 x 160 μm, or a cross-sectional area of ​​20,000 to 25,000 μm². As shown below, LLA content exceeding 1 wt% or 2 wt% is likely to reduce the time to maintain radial and mechanical strength necessary to successfully cure the affected segment.

The LLA monomer can be dispersed in the form of powder or solid particles throughout the body or part of the stent. These particles may have a size below 100 nm, from 100 nm to 1000 nm or above 1000 nm, the size being the diameter or other characteristic length. Optionally, the LLA monomer can be mixed or dissolved at the molecular level with PLLA.

In in vitro and in vivo studies described below on the degradation of stents from PLLA containing L-lactide monomer, LLA has been shown to dramatically and unexpectedly accelerate the rate of stent degradation, especially when its content exceeds 1 wt%. Stents containing LLA monomer at a mass percentage greater than 1 wt% lose mechanical integrity and erode too quickly to allow successful treatment for implantation in coronary or peripheral vessels. In addition, low LLA concentration is an advantage because the dispersed monomer in the polymer has no or little effect on the mechanical properties of the polymer.

In addition, uniform or largely uniform dispersion of the LLA monomer in PLLA is important to ensure a uniform degradation pattern of the entire stent body. It has been shown that at such low concentrations of LLA monomer, the uniformity of distribution largely depends on how the LLA is mixed or dispersed. Accordingly, additional embodiments include the LLA blending method of the PLLA polymer from which the stent is made.

The results of the in vitro and in vivo degradation scheme can be used to assess the effect of LLA monomer concentration on the stent degradation scheme from PLLA. In addition, this impact can be assessed on the basis of theoretical models.

The model of hydrolytic degradation of aliphatic polyesters is expressed by the formula Mn (t) = Mn (0) exp (-Kt), where Mn (t) is the number average molecular weight at t, Mn (0) is the number average molecular weight at t = 0 and K is the hydrolytic degradation rate constant. Pitt, CG, J. of Applied Polymer Science 26, 3779-3787 (1981); Pitt, CG, Biomaterials 2, 215-220 (1981); Weir, NA, Proceedings of the Institution of Mechanical Engineers, Part H: J. of Engineering in Medicine 218, 307-319 (2004); Weir, NA, Part H: J. of Engineering in Medicine 218, 321-330 (2004). Model assumptions are rational provided that there is no weight loss, because the weight loss would have an adverse effect on the concentration of water and carboxyl end groups in the sample. This formula can also be written as follows. ln [Mn (t) / Mn (0)] = Kt. Thus, by comparing the data for Mn (t) / Mn (0) versus t on the log-line graph, the hydrolytic degradation rate can be determined based on the slope of the line connecting the points.

In vitro and in vivo data on degradation of PLLA stents containing LLA monomer at various concentrations was collected to assess the effect of LLA concentration on the PLLA degradation scheme. In vivo data was obtained using an experimental animal model.

Parameters and properties such as Mn, radial strength, appearance of stent cracks or fractures (structural / mechanical integrity) and degradation time were monitored. In all tests, the stent underwent the treatment described in this document. The parameters of these stents are shown in Table 1 below. The stents were covered with poly (DLlactide) and everolimus coating applied to the scaffolding.

Table 1. Summary of stent parameters in degradation studies

<td>Technical specifications</td><td>Value</td>

<td>Skeletal polymer (PLLA)</td><td></td>

<td>Molecular mass (Mw)</td><td> 180000-200000</td>

<td>Mn</td><td> 90000-100000</td>

<td>Stent weight (length 18 mm)</td><td>9.0 mg</td>

<td>Mass to unit of length ratio</td><td>0.5 mg / mm</td>

<td>crystallinity</td><td>45% (determined by differential scanning calorimetry [DSC])</td>

<td>Cross-section of the expansion element (bracket)</td><td>150 microns x 150 microns</td>

<td>Coating thickness</td><td>3 microns</td>

<td>Coating weight</td><td>308 μg (polymer: everolimus ratio 1: 1)</td>

<td>Coating polymer</td><td></td>

<td>Molecular mass (Mw)</td><td> 66000</td>

<td>Mn</td><td> 39000</td>

<td>Blow molding</td><td></td>

<td>Percentage expansion in the radial plane</td><td> 400%</td>

<td>Percentage elongation in the axial plane</td><td> 20%</td>

<td>Mechanical laser processing</td><td>Femtosecond laser (120 fs)</td>

In vitro tests were performed using a phosphate buffer salt solution (PBS buffer) at 37 ° C. The Mn value of the stent was measured by gel permeation chromatography (GPC) using polystyrene standards. The animal model for in vivo testing was the Yucatan miniature pig.

In the in vitro study, a plot of Mn versus time was made as a function of ln [Mn (t) / Mn (0)] versus time to assess the predictive value of the hydrolytic degradation model. This model was then used to assess the effect of LLA on PLLA degradation.

In vitro studies also measured the Mn value as a function of time for PLLA stents containing LLA monomer at various concentrations. In these studies, LLA monomer was added to the PLLA resin as described herein. FIG. 2 is a graph of ln [Mn (t) / Mn (0)] versus time for four groups of stents: nominal (0.05 wt% LLA) and containing 0.2 wt% LLA, 0.55 wt% LLA, 1 wt% LLA and 3.3 wt% LLA in PLLA.

Data for each concentration is represented as a straight line, the slope of which indicates the rate constant (K) of the degradation model. The K rate constants for each LLA monomer concentration are summarized in Table 2 and FIG. 3.

Table 2. K rate constant for the degradation model from in vitro data

<td>Stent group: wt% LLA</td><td>k (x 10³) (days^-1)</td><td>R2</td>

<td>nominal</td><td> 1.9</td><td> 0.962</td>

<td> 0.2</td><td> 3.0</td><td> 0.972</td>

<td> 0.55</td><td> 7.2</td><td> 0.969</td>

<td> 1</td><td> 13.4</td><td> 0.960</td>

<td> 3.3</td><td> 48.8</td><td> 0.989</td>

The R2 determination coefficient is given for each data group to assess the predictability of the linear degradation model. The closer the R2 value is to 1, the more reliable the predictability of the model. The data in Table 2 show the surprisingly radical impact of LLA monomer content on stent degradation. For example, when the LLA concentration increases from a value close to 0 to

0.55 wt%, the K constant increases almost threefold, and when the LLA concentration increases from a value close to 0 to wt%, the speed constant increases almost sixfold. The difference in molecular weight change over time as the LLA concentration increases is so radical. Table 3 shows the percentage reduction in Mn and Mn at time points based on model predictions for an LLA content of 1 wt% for the second data set.

Table 3. Molecular weight change over time for PLLA with 1 wt% LLA

<td>Days</td><td>Decline</td><td>Mn</td>

<td> 20</td><td> 23%</td><td>77 k</td>

<td> 40</td><td> 41%</td><td>59 k</td>

<td> 60</td><td> 55%</td><td>45 k</td>

<td> 80</td><td> 65%</td><td>35 k</td>

<td> 120</td><td> 80%</td><td>20 k</td>

As shown above, mechanical integrity begins to decrease before weight loss. In addition, radial strength decreases before or in parallel with a weakening of mechanical integrity. When the molecular weight of crosslinking is about 17,000, the polymer loses its previous mechanical properties and breaks down under load. The mechanical integrity of the stent disappears well before reaching the molecular weight of the crosslinking.

As shown above, the onset of the in vitro mass loss process is probably the upper limit for in vivo degradation, which means that in vivo weight loss occurs earlier. In addition, mechanical integrity and radial strength may also begin to decline earlier in vivo rather than in vitro. Table 4 provides in vivo and in vitro weight loss data for PLLA stents containing no LLA monomers. Both the time difference between the start of the weight loss process and the degree of weight loss is significant.

Table 4. Comparison of in vitro and in vivo weight loss for a stent with PLLA without LLA monomers

<td>Time (months)</td><td>In vitro weight loss</td><td>In vivo weight loss</td>

<td> 9</td><td> 0%</td><td> 20%</td>

<td> 12</td><td> 0%</td><td> 30%</td>

<td> 18</td><td> -</td><td> 35%</td>

<td> 21</td><td> 9%</td><td> -</td>

It is believed that in the case of implantation in coronary and peripheral arteries, mechanical integrity should not change (no significant fractures, e.g. strut elements) for at least 2 to 4 months after implantation, i.e. in time allowing the stent to be embedded in the vessel wall. In addition, it is believed that radial strength should not change for at least about 1 month to avoid unfavorable reconstruction of the vessel wall. It is anticipated that radial strength will decline before a loss of mechanical integrity occurs, and mechanical integrity begins to decrease before the weight loss process begins. An introduction to the loss of mechanical integrity is the formation of cracks in the stent. Thus, based on in vitro data, it was found that for coronary and peripheral vessels implantation, the LLA concentration in the stent should be less than 1 wt% in order to maintain radial strength and mechanical integrity at the desired time.

Other in vitro studies evaluated stent groups with LLA content in PLLA of 0 wt%, 0.2 wt%, 0.55 wt%, 1 wt% and approximately 3 wt%, which were observed for four months. Radial strength and stent integrity were monitored in these studies. Table 5 summarizes the observed changes in radial strength and mechanical integrity demonstrated in studies of radial strength of various groups of stents. As shown in Table 5, for LLA content close to 0 wt%, 0.2 wt% and 0.55 wt%, the radial strength is maintained for a maximum of over 4 months and no fractures were observed during this period. For an LLA content of 1 wt% LLA, the radial strength decreases constantly over a period of about 1 1 month to about 3 months. Radial strength may decrease earlier in vivo. In addition, the appearance of significant fractures after just 42 days indicates an early loss of radial strength and mechanical integrity. The above results indicate that a stent with an LLA content exceeding 1 wt% is probably not suitable for implantation in coronary or peripheral vessels. A decrease in radial strength and the appearance of significant fractures in a stent containing about 3 wt% LLA indicates that a PLLA stent with this LLA content is completely unsuitable.

Table 5. Radial strength and mechanical integrity of PLLA stents with different content of LLA monomers

<td>LLA content (wt%)</td><td>Radial strength</td><td>Mechanical integrity</td>

<td> 0.05</td><td>retained for a maximum of over 4 months (126 days)</td><td>no fractures for a maximum of over 4 months (126 days)</td>

<td> 0.2</td><td>retained for a maximum of over 4 months (126 days)</td><td>no fractures for a maximum of over 4 months (126 days)</td>

<td> 0.55</td><td>retained for a maximum of over 4 months (126 days)</td><td>no fractures for a maximum of over 4 months (126 days)</td>

<td> 1.0</td><td>it steadily decreases from ~ 1% of the month to about 3 months</td><td>fragmentation of stent segments visible after 42 days</td>

<td> ~ 3</td><td>begins to decrease steadily after about 2 weeks</td><td>fragmentation of stent segments visible after 14 days</td>

Pre-clinical in vivo (animal) data was collected for stent groups with LLA content of 0 wt%, 0.1 wt%, 0.4 wt%, approximately 0.6 wt%, 1 wt% and 3.8 wt% obtained within a maximum of 28 days after implantation . For stents with an LLA content of 0.4 wt% and about 0.6 wt%, no fractures were found for 28 days after implantation. For stents with an LLA content of 1 wt%, fractures were demonstrated 28 days after implantation. In the case of stents with an LLA content of 3.8 wt%, significant fractures were observed already after 7 days, and the stent broke to pieces after 28 days.

The inventor has found that by forming PLLA tubes with a predicted LLA content of 0.05 - 0.5 wt% by mechanically mixing the required amount of LLA in PLLA, it is not possible to obtain a product in which the LLA monomer would be mixed evenly. Significant fluctuations in LLA concentration were observed in stents made of tubes formed in this way.

Two methods are presented for even mixing of components and obtaining a constant concentration of LLA monomer in various tubes and stents made of this material.

The first method involves making a concentrate of a mixture of PLLA and LLA with a LLA concentration higher than the target concentration. The concentrate is prepared by dissolving LLA and PLLA in a solvent such as chloroform. By evaporating chloroform, a concentrate is obtained which is a homogeneous mixture of PLLA and LLA. Then the concentrate is mixed in the smelting process (e.g. in an extruder) with PLLA in an amount sufficient to achieve the target LLA concentration. The following example illustrates this method:

Step 1: Dissolve 2 g LLA monomer and 8 g PLLA in 400 ml chloroform.

Step 2: Evaporate the chloroform to obtain a homogeneous mixture of PLLA and LLA with an LLA content of 25 wt%.

Step 3: Mix in the extruder a mixture containing 25 wt% LLA with 4 kg PLLA to obtain a PLLA with an LLA content of 0.5% wt%.

The second method for preparing a homogeneous LLA mixture in PLLA is to dissolve the LLA in a solvent such as chloroform to obtain a solution and spray it on the PLLA granulate.

Then the solvent is removed and the LLA remains deposited on the PLLA granulate. PLLA granules are smelted in an extruder and formed into tubes containing a target concentration of LLA in PLLA. The following example illustrates this method:

Step 1: Dissolve 0.5 g LLA in 100 mL anhydrous methanol to make a solution.

Stage 2: Spray the solution onto 1 kg of PLLA granules and mix.

Step 3: Place the granulate in a vacuum oven to remove the solvent.

Step 4: Place the granulate in an extruder and form tubes with an LLA content of 0.5 wt%.

Proxy:

"ATENTOWA" BELLEPAT "LAW OFFICE

Izabeia Szychulska-Hawranek ul Słowackiego 44, 37-700 Ρΐ2 * · ιΐνέΙ phone (016) 7o2-37-77 fax: (016) 375-02-87 phone, mobile, (0608) 503-081 e-mail <a href="mailto:bellepat@op.pl">bellepat@op.pl</a> NIP (tax identification number): 795-207-16-72 REGON (tax identification number): 1803505 (6

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