Biodegradable stent with adjustable degradation rate

BACKGROUND OF THE INVENTION

Field of the Invention

0001This invention relates to methods of treatment of blood vessels with bioabsorbable polymeric medical devices, in particular, stents.

Description of the State of the Art

0002This invention relates to radially expandable endoprostheses, that are adapted to be implanted in a bodily lumen. An "endoprosthesis" corresponds to an artificial device that is placed inside the body. A "lumen" refers to a cavity of a tubular organ such as a blood vessel. A stent is an example of such an endoprosthesis. Stents are generally cylindrically shaped devices that function to hold open and sometimes expand a segment of a blood vessel or other anatomical lumen such as urinary tracts and bile ducts. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. "Stenosis" refers to a narrowing or constriction of a bodily passage or orifice. In such treatments, stents reinforce body vessels and prevent restenosis following angioplasty in the vascular system. "Restenosis" refers to the reoccurrence of stenosis in a blood vessel or heart valve after it has been treated (as by balloon angioplasty, stenting, or valvuloplasty) with apparent success.

0003Stents are typically composed of scaffolding that includes a pattern or network of interconnecting structural elements or struts, formed from wires, tubes, or sheets of material rolled into a cylindrical shape. This scaffolding gets its name because it physically holds open and, if desired, expands the wall of the passageway. Typically, stents are capable of being compressed or crimped onto a catheter so that they can be delivered to and deployed at a treatment site.

0004Delivery includes inserting the stent through small lumens using a catheter and transporting it to the treatment site. Deployment includes expanding the stent to a larger diameter once it is at the desired location. Mechanical intervention with stents has reduced the rate of restenosis as compared to balloon angioplasty. Yet, restenosis remains a significant problem. When restenosis does occur in the stented segment, its treatment can be challenging, as clinical options are more limited than for those lesions that were treated solely with a balloon.

0005Stents are used not only for mechanical intervention but also as vehicles for providing biological therapy. Biological therapy uses medicated stents to locally administer a therapeutic substance. The therapeutic substance can also mitigate an adverse biological response to the presence of the stent. Effective concentrations at the treated site require systemic drug administration which often produces adverse or even toxic side effects. Local delivery is a preferred treatment method because it administers smaller total medication levels than systemic methods, but concentrates the drug at a specific site. Local delivery thus produces fewer side effects and achieves better results.

0006A medicated stent may be fabricated by coating the surface of either a metallic or polymeric scaffolding with a polymeric carrier that includes an active or bioactive agent or drug. Polymeric scaffolding may also serve as a carrier of an active agent or drug.

0007The stent must be able to satisfy a number of mechanical requirements. The stent must be have sufficient radial strength so that it is capable of withstanding the structural loads, namely radial compressive forces, imposed on the stent as it supports the walls of a vessel. Once expanded, the stent must adequately provide lumen support during a time required for treatment in spite of the various forces that may come to bear on it, including the cyclic loading induced by the beating heart. In addition, the stent must possess sufficient flexibility with a certain resistance to fracture.

0008Stents made from biostable or non-erodible materials, such as metals, have become the standard of care for percutaneous coronary intervention (PCI) as well as in peripheral applications, such as the superficial femoral artery (SFA), since such stents have been shown to be capable of preventing early and later recoil and restenosis.

0009In order to effect healing of a diseased blood vessel, the presence of the stent is necessary only for a limited period of time. There are certain disadvantages to the presence of a permanent implant in a vessel such as compliance mismatch between the stent and vessel and risk of embolic events. To alleviate such disadvantages, stent can be made from materials that erode or disintegrate through exposure to conditions within the body. Thus, erodible portions of the stent can disappear from the implant region after the treatment is completed, leaving a healed vessel. Stents fabricated from biodegradable, bioabsorbable, and/or bioerodable materials such as bioabsorbable polymers can be designed to completely erode only after the clinical need for them has ended.

0010Like a durable stent, a biodegradable stent must meet time dependent mechanical requirements. For example, it must provide patency for a minimum time period. However, it is also important for a biodegradable stent to completely degrade from the implant site with in a certain period of time. A biodegradable material that can provide the mechanical requirements may not possess the requisite or desired , degradation time. In addition, the requisite or desired degradation time varies between types of applications, i.e. coronary or peripheral.

0011<patcit id="pcit0001" dnum="EP1184008A1"><text>EP 1 184 008 A1</text></patcit> discloses: stent for treating a diseased section of a blood vessel, comprising: a polymeric scaffolding consisting of poly(L-lactide), wherein the crystallinity of the poly(L-lactide) is between 15-60%, and wherein the weight average molecular weight of the poly(L-lactide) is between 60,000 and 300,000, relative to polystyrene standards.

0012<patcit id="pcit0002" dnum="US2007253999A"><text>US 2007/253999</text></patcit> and <patcit id="pcit0003" dnum="WO2007146354A2"><text>WO 2007/146354 A2</text></patcit> disclose biodegradable poly (Llactide) stents.

0013<nplcit id="ncit0001" npl-type="s"><text>BENDIX D: "Chemical synthesis of polylactide and its copolymers for medical applications", POLYMER DEGRADATION AND STABILITY, BARKING, GB</text></nplcit>, discloses that monomers of L-lactide in PLLA are considered impurities that lead to severe (rapid) degradation during thermal processing.

SUMMARY OF THE INVENTION

0014Various embodiments of the present invention include a stent for treating a diseased section of a blood vessel, comprising: a polymeric scaffolding comprising poly(L-lactide) and between 0.05 wt% and 0.5 wt% L-lactide monomer mixed, dispersed, or dissolved within the poly(L-lactide), wherein the crystallinity of the poly(L-lactide) is between 20-50%, and wherein the number average molecular weight of the poly(L-lactide) is between 60,000 and 300,000, relative to polystyrene standards; and wherein the scaffolding comprises greater than 95wt% poly(L-lactide).

0015It is disclosed a method of treating a diseased section of a blood vessel, comprising: deploying a bioabsorbable polymeric stent at a diseased section of a blood vessel the stent comprising a body including a scaffolding composed of a pattern of struts, wherein the body is composed of poly(L-lactide) and between 0.05 and 0.5 % L-lactide monomer, wherein the stent supports the vessel wall at or near a deployment diameter for a period of support which is followed by a decline in radial strength of the stent such that the stent is unable to support the vessel wall, wherein the scaffolding breaks apart and is completely absorbed away into the human body.

BRIEF DESCRIPTION OF THE DRAWINGS

0016<ul id="ul0001" list-style="none" compact="compact"><li><figref idref="f0001">FIG. 1</figref> depicts an exemplary stent.</li><li><figref idref="f0002">FIG. 2</figref> depicts a plot of the <i>in vitro</i> degradation behavior for a PLLA stent with different concentrations of L-lactide monomer.</li><li><figref idref="f0003">FIG. 3</figref> depicts a plot of a degradation rate constants versus L-lactide monomer concentration based on the data of <figref idref="f0002">FIG. 2</figref>.</li></ul>

DETAILED DESCRIPTION OF THE INVENTION

0017Coronary arteries refer generally to arteries that branch off the aorta to supply the heart muscle with oxygenated blood. Peripheral arteries refer generally to blood vessels outside the heart and brain.

0018In both coronary artery disease and peripheral artery disease, the arteries become hardened and narrowed or stenotic and restrict blood flow. In the case of the coronary arteries, blood flow is restricted to the heart, while in the peripheral arteries blood flow is restricted leading to the kidneys, stomach, arms, legs, and feet. The narrowing is caused by a the buildup of cholesterol and other material, called plaque, on the inner walls of the vessel. Such narrowed or stenotic portions are often referred to as lesions. Artery disease also includes the reoccurrence of stenosis or restenosis that occurs after an angioplasty treatment. Although there are probably several mechanisms that lead to restenosis of arteries, an important one is the inflammatory response, which induces tissue proliferation around an angioplasty site. The inflammatory response can be caused by the balloon expansion used to open the vessel, or if a stent is placed, by the foreign material of the stent itself.

0019Embodiments of the present invention are applicable to treatment of coronary and peripheral disease in coronary arteries and various peripheral vessels including the superficial femoral artery, the iliac artery, and carotid artery. The embodiments are further applicable to various stent types, such as self-expandable and balloon

0020The stent is made from an extruded PLLA polymer tube that has been radially expanded and axially stretched to provide the induced orientation. The polymer tube is expanded by blow molding with a percent radial expansion between 200% and 500%, and a percent axial stretch from 20% to 200%. The extruded PLLA tubing has a percent of axial stretch from 100% to 400%. The stent is formed from the expanded tube by laser cutting the tubing in its expanded state.

0021Additionally, the blow molding process is performed in a manner that result in small crystallites dispersed through an amorphous matrix. Prior to expansion, the tube is heated to a temperature between 65 °C and 75 °C to induce formation of smaller crystallites that enhance fracture toughness. The tube is quenched below the glass transition temperature (Tg) after expansion to prevent further crystal growth. The degree of crystallinity is 20-50%. Below 20% crystallinity the stent body may not be have sufficient strength, while above 50% crystallinity the stent body may be too brittle. The number average molecular weight (Mn) of scaffolding material in the final product (in g/mol) is between 60,000 and 300,000, or more narrowly between 80,000 and 200,000.

0022An exemplary strut cross-section can be rectangular, for example, 140 x 140 µm to 160 x 160 µm or having a cross-sectional area between 20,000 and 25,000 µm².

0023As shown below, it is expected that an LLA content of greater than 1 wt% or greater than 2 wt% will result in a stent that does not maintain radial strength and mechanical strength long enough for effective treatment of a diseased section of a vessel.

0024The LLA monomer can be dispersed in the form of a powder or particular particles through all or a portion of a stent body. The size of such particles can be less than 100 nm, between 100 nm and 1000 nm, or greater than 1000 nm, where size can refer to diameter or some other characteristic length. Alternatively, the LLA monomer can be mixed or dissolved on a molecular level with the PLLA.

0025It has been observed from <i>in vitro</i> and <i>in vivo</i> degradation studies of PLLA stents with L-lactide monomer, discussed below, that LLA provides a dramatic and unexpected increase in the degradation rate of the stent, particularly above about 1 wt%. Stents having monomer compositions above about 1 wt% LLA lose mechanical strength, lose mechanical integrity, and erode away too fast for effective treatment in coronary and peripheral applications. Additionally, the low concentration of LLA are advantageous since the effect of the dispersed monomer in the polymer has no or a minimal effect on the mechanical properties of the polymer.

0026Additionally, it is important that the LLA monomer be uniformly or substantially uniformly dispersed through the PLLA of the stent to provide uniform degradation behavior through the body of the stent. It has been observed that for such low concentrations of LLA monomer, the uniformity of distribution is highly dependent on the manner of mixing or dispersing of the LLA. Thus, additional embodiments include a method of mixing the LLA in the PLLA stent material.

0027<i>In vitro</i> and <i>in vivo</i> studies of degradation behavior can be used to assess the influence of LLA monomer concentration on the degradation behavior of a PLLA stent. In addition, the influence can be assessed using theoretical models.

0028A hydrolytic degradation model for aliphatic polyesters having the form Mn(t) = Mn(0)exp(-Kt), wherein Mn(t) is the number average molecular weight at time t, Mn(0) is the number average molecular weight at t=0, and K is the hydrolytic degradation rate constant. <nplcit id="ncit0002" npl-type="s"><text>Pitt, C.G., J. of Applied Polymer Science 26, 3779-3787 (1981</text></nplcit>); <nplcit id="ncit0003" npl-type="s"><text>Pitt, C.G., Biomaterials 2, 215-220 (1981</text></nplcit>); <nplcit id="ncit0004" npl-type="s"><text>Weir, N.A., Proceedings of the Institution of Mechanical Engineers, Part H: J. of Engineering in Medicine 218, 307-319 (2004</text></nplcit>); <nplcit id="ncit0005" npl-type="s"><text>Weir, N.A., Part H: J. of Engineering in Medicine 218, 321-330 (2004</text></nplcit>). The assumptions inherent in the model are reasonable provided that the mass loss has not occurred, since mass loss would affect the concentrations of water and carboxylic end groups in the sample. The equation can also be written as: In [Mn(t)/Mn(0)] = - Kt. Therefore, by representing data for Mn(t)lMn(0) versus t on a log-linear plot, one may infer the hydrolytic degradation rate from the slope of the connecting points.

0029<i>In vitro</i> and <i>in vivo</i> degradation data has been generated for PLLA stents with different concentrations of LLA monomer to investigate its affect on the degradation behavior of a PLLA stent. The <i>in vivo</i> data was obtained using an animal model. Parameters and behavior that have been monitored include Mn, radial strength, appearance of cracks or fractures in stent struts (structural/mechanical integrity), and degradation time. In all studies, the stent is processed as described herein. The parameters of the stents are shown in Table 1 below. The stents included a coating over a scaffolding composed of poly(DL-lactide) and Everolimus. <tables id="tabl0001" num="0001"><table frame="all"><title>Table 1. Summary stent parameters in degradation studies</title><tgroup cols="2"><colspec colnum="1" colname="col1" colwidth="43mm" /><colspec colnum="2" colname="col2" colwidth="52mm" /><thead><row><entry align="center" valign="top">Specification</entry><entry align="center" valign="top">Value</entry></row></thead><tbody><row><entry>Backbone polymer (PLLA)</entry><entry /></row><row><entry align="center">Mw</entry><entry>180,000-200,000</entry></row><row><entry align="center">Mn</entry><entry>90,000-100,000</entry></row><row><entry>Mass of stent (18mm length)</entry><entry>9.0 mg</entry></row><row><entry>Mass / unit length</entry><entry>0.5 mg /mm</entry></row><row><entry>Crystallinity</entry><entry>45 % (as measured by DSC)</entry></row><row><entry>Strut cross section</entry><entry>150 micron x 150 micron</entry></row><row><entry>Coating thickness</entry><entry>3 microns</entry></row><row><entry>Coating mass</entry><entry>308 µg (1:1 polymer:Everolimus)</entry></row><row><entry>Coating polymer</entry><entry /></row><row><entry align="center">Mw</entry><entry>66,000</entry></row><row><entry align="center">Mn</entry><entry>39,000</entry></row><row><entry>Blow molding:</entry><entry /></row><row><entry align="center">Percent radial expansion</entry><entry>400%</entry></row><row><entry align="center">Percent axial elongation</entry><entry>20%</entry></row><row><entry>Laser machining</entry><entry>120 fs laser</entry></row></tbody></tgroup></table></tables>

0030The <i>in vitro</i> studies were performed in a phosphate buffered saline (PBS buffer) solution at 37 °C. The Mn of the stent was measured by GPC using polystyrene standards. The animal model for the <i>in vivo</i> studies was_Yucatan mini swine_.

0031In the <i>in vitro</i> study, the Mn dependence on time is plotted in the form of ln[Mn(t)/Mn(0)] versus time to assess the predictive ability of the hydrolytic degradation model. The model is then used to assess the effect of LLA on the degradation of PLLA.

0032In the <i>in vitro</i> studies, the Mn as a function of time was also measured for PLLA stents having different concentrations of LLA monomer. In these studies, the LLA monomer was added to the PLLA resin in a manner described herein. <figref idref="f0002">FIG. 2</figref> depicts a plot of ln[Mn(t)/Mn(0)] versus time for this set for four groups of stents: nominal (0.05 wt % LLA), 0.2 wt% LLA, 0.55 wt% LLA, 1 wt% LLA, and 3.3 wt% LLA monomer in PLLA.

0033The data for each concentration is fitted with a straight line, the slope of which provides the rate constant, K, of the degradation model. The rate constant, K, for each concentration of LLA monomer is summarized in Table 2 and <figref idref="f0003">FIG. 3</figref>. <tables id="tabl0002" num="0002"><table frame="all"><title>Table 2. Rate constant K for degradation model from <i>in vitro</i> data.</title><tgroup cols="3"><colspec colnum="1" colname="col1" colwidth="36mm" /><colspec colnum="2" colname="col2" colwidth="32mm" /><colspec colnum="3" colname="col3" colwidth="29mm" /><thead><row><entry align="center" valign="middle">Stent Group: wt% LLA</entry><entry align="center" valign="middle"><i>k</i> (x 10³) (days^-1)</entry><entry align="center" valign="middle"><i>R</i>2</entry></row></thead><tbody><row><entry align="center" valign="middle">Nominal</entry><entry align="center" valign="middle">1.9</entry><entry align="center" valign="middle">0.962</entry></row><row><entry align="center" valign="middle">0.2</entry><entry align="center" valign="middle">3.0</entry><entry align="center" valign="middle">0.972</entry></row><row><entry align="center" valign="middle">0.55</entry><entry align="center" valign="middle">7.2</entry><entry align="center" valign="middle">0.969</entry></row><row><entry align="center" valign="middle">1</entry><entry align="center" valign="middle">13.4</entry><entry align="center" valign="middle">0.960</entry></row><row><entry align="center" valign="middle">3.3</entry><entry align="center" valign="middle">48.8</entry><entry align="center" valign="middle">0.989</entry></row></tbody></tgroup></table></tables> The coefficient of determination, R2, is given for each group of data to assess the predictability of the linear degradation model. The closer R2 is to 1, the more reliable the predictive ability of the model. The data in Table 2 shows an unexpectedly dramatic effect of LLA monomer content on the degradation of the stent. For example, when the LLA concentration increases from near 0 to 0.55 wt%, K increases near three times and from near 0 to 1 wt%, the rate constant increases near six times. The differences in the change in the molecular weight with time as the LLA concentration increases is just as dramatic. Table 3 shows the percent drop in Mn and the Mn at time points based on the model predictions for 1 wt% LLA for the second set of data. <tables id="tabl0003" num="0003"><table frame="all"><title>Table 3. Change in molecular weight with time for PLLA with 1 wt% PLLA</title><tgroup cols="3"><colspec colnum="1" colname="col1" colwidth="36mm" /><colspec colnum="2" colname="col2" colwidth="35mm" /><colspec colnum="3" colname="col3" colwidth="33mm" /><thead><row><entry align="center" valign="top">Days</entry><entry align="center" valign="top">Drop in Mn</entry><entry align="center" valign="top">Mn</entry></row></thead><tbody><row><entry align="center">20</entry><entry align="center">23%</entry><entry align="center">77 k</entry></row><row><entry align="center">40</entry><entry align="center">41%</entry><entry align="center">59 k</entry></row><row><entry align="center">60</entry><entry align="center">55%</entry><entry align="center">45 k</entry></row><row><entry align="center">80</entry><entry align="center">65%</entry><entry align="center">35 k</entry></row><row><entry align="center">120</entry><entry align="center">80%</entry><entry align="center">20 k</entry></row></tbody></tgroup></table></tables> As indicated above, the loss of mechanical integrity starts before mass loss. In addition, the loss of radial strength occurs before or along with loss of mechanical integrity. At the entanglement molecular weight, about 17,000, the polymer no longer has mechanical properties and disintegrates under a applied load. The stent loses mechanical integrity well before the entanglement molecular weight.

0034As indicated above, the <i>in vitro</i> start of mass loss is expected to be an upper limit for <i>in vivo</i> degradation, that is, mass loss starts at an earlier time <i>in vivo.</i> In addition, the start of mechanical integrity loss and the loss of radial strength might also occur at an earlier time <i>in vivo</i> than <i>in vitro.</i> Table 4 shows <i>in vivo</i> and <i>in vitro</i> data for mass loss for a PLLA stent with no LLA monomer. The difference in the time that mass loss starts and the difference in the degree of loss is significant. <tables id="tabl0004" num="0004"><table frame="all"><title>Table 4. Comparison of <i>in vitro</i> and <i>in vivo</i> mass loss for PLLA stent with no LLA monomer.</title><tgroup cols="3"><colspec colnum="1" colname="col1" colwidth="43mm" /><colspec colnum="2" colname="col2" colwidth="42mm" /><colspec colnum="3" colname="col3" colwidth="42mm" /><thead><row><entry align="center" valign="top">Time (months)</entry><entry align="center" valign="top"><i>In vitro</i> mass loss</entry><entry align="center" valign="top"><i>In vivo</i> Mass loss</entry></row></thead><tbody><row><entry align="center">9</entry><entry align="center">0%</entry><entry align="center">20%</entry></row><row><entry align="center">12</entry><entry align="center">0%</entry><entry align="center">30%</entry></row><row><entry align="center">18</entry><entry align="center">-</entry><entry align="center">35%</entry></row><row><entry align="center">21</entry><entry align="center">9%</entry><entry align="center">-</entry></row></tbody></tgroup></table></tables>

0035For coronary and peripheral applications, it is believed that the mechanical integrity should remain intact for at least 2 to 4 months without severe fractures (e.g. breaking of struts) after implantation to allow incorporation of stent into vessel wall. Additionally, it is believed that radial strength should be maintained for at least about 1 month to prevent negative remodeling. The radial strength is expected to be lost prior to the mechanical integrity and the start of the loss of mechanical integrity is expected to start before mass loss. A prelude to the start of the loss of mechanical integrity corresponds to the formation of cracks in the stent. Therefore, based on this <i>in vitro</i> data, the LLA concentration should be less than 1 wt% for the stent to maintain radial strength and mechanical integrity for the desired times for coronary and peripheral applications.

0036In another set of <i>in vitro</i> studies, stent groups with near 0 wt%, 0.2 wt%, 0.55 wt%, 1 wt% and about 3 wt% of LLA in PLLA were studied and tracked over a four month period. In these studies, the radial strength and stent integrity were tracked. Table 5 summarizes the observed change in radial strength and mechanical integrity after radial strength testing for the different groups. As shown in Table 5, for near 0 wt%, 0.2 wt%, and 0.55 wt% LLA, the radial strength is maintained up to more than 4 months and there were no fractures observed up to more than 4 months. For 1 wt% LLA, the radial strength showed a consistent decrease between about 1 ½ months to about 3 months. The drop in radial strength might occur sooner <i>in vivo.</i> Additionally, the significant fractures at only 42 days indicate an early loss of radial strength and mechanical integrity. These results indicate a stent with greater than 1 wt% LLA is very likely unsuitable for coronary or peripheral applications. The drop in radial strength and significant fractures in the stent with about 3 wt% LLA indicate that a PLLA stent with this concentration is totally unsuitable. <tables id="tabl0005" num="0005"><table frame="all"><title>Table 5. Radial strength and mechanical integrity of PLLA stents with different LLA monomer content.</title><tgroup cols="3"><colspec colnum="1" colname="col1" colwidth="37mm" /><colspec colnum="2" colname="col2" colwidth="66mm" /><colspec colnum="3" colname="col3" colwidth="63mm" /><thead><row><entry align="center" valign="top"><b>LLA content (wt%)</b></entry><entry align="center" valign="top"><b>Radial strength</b></entry><entry align="center" valign="top"><b>Mechanical integrity</b></entry></row></thead><tbody><row><entry align="center">0.05</entry><entry>maintained up to more than 4 months (126 days)</entry><entry>no fractures at up to more than 4 months (126 days)</entry></row><row><entry align="center">0.2</entry><entry>maintained up to more than 4 months (126 days)</entry><entry>no fractures at up to more than 4 months (126 days)</entry></row><row><entry align="center">0.55</entry><entry>maintained up to more than 4 months (126 days)</entry><entry>no fractures at up to more than 4 months (126 days)</entry></row><row><entry align="center">1.0</entry><entry>consistent decrease between ∼1 ½ months to about 3 months</entry><entry>stent segments in pieces observed at 42 days</entry></row><row><entry align="center">∼ 3</entry><entry>consistent decrease starting at about 2 weeks</entry><entry>stent segments in pieces observed at 14 days</entry></row></tbody></tgroup></table></tables>

0037Preclinical <i>in vivo</i> (animal) results for stent groups with a 0 wt%, 0.1 wt%, 0.4 wt % LLA, ca. 0.6 wt% LLA, 1 wt% LLA, and 3.8 wt% LLA have been obtained for up to 28 days after implantation. For stents with 0.4 wt% LLA and ca. 0.6% there were no fractures observed at 28 days after implantation. For stents with 1 wt% LLA, fractures were observed at 28 days after implantation. For stents with 3.8 wt% LLA, there were significant fractures at only 7 days and the stent broke into pieces after 28 days.

0038It has been observed by the inventor that forming a PLLA tube intended to have between 0.05 - 0.5 wt% LLA by mechanical blending of the requisite amount of LLA in PLLA fails to result in tubes with uniform blending of the LLA. Large variations in LLA concentration between stents formed from tubes in this manner were observed.

0039Two methods are provided that result in uniform blending with consistent concentration in LLA monomer in different tubes and stents fabricated therefrom. The first method includes making a master batch of PLLA and LLA mixture with concentrations of LLA larger than the target concentration. The master batch is made by dissolving LLA and PLLA in a solvent such as chloroform. The chloroform is evaporated to form the master batch with is a uniform mixture of PLLA and LLA. The master batch is then mixed by melt processing, such as in an extruder, with a sufficient amount of PLLA to obtain the target LLA concentration. The method is illustrated by the following example: <ul id="ul0002" list-style="none" compact="compact"><li>Step 1: Dissolve 2g LLA monomer and 8 g PLLA in 400 ml chloroform</li><li>Step 2: Evaporate chloroform to form uniform mixture of PLLA and LLA with 25 wt% LLA.</li><li>Step 3: Blend 25 wt% LLA mixture in extruder with 4 kg of PLLA to obtain 0.5 wt% LLA in PLLA.</li></ul> The second method of preparing a uniform mixture of LLA in PLLA includes dissolving LLA in a solvent such as chloroform to form a solution and spraying the solution on PLLA pellets. The solvent is removed leaving the LLA deposited on the PLLA pellets. The PLLA pellets are melt processed, in an extruder, for example, to form tubes with the target concentration of LLA in PLLA. The method is illustrated by the following example: <ul id="ul0003" list-style="none" compact="compact"><li>Step 1: Dissolve 0.5 g LLA in 100 ml anhydrous methanol to form solution.</li><li>Step 2: Spray solution onto 1 kg of PLLA pellets and stir.</li><li>Step 3: Place pellets in vacuum oven to remove solvent.</li><li>Step 4: Place pellets in extruder and form tube with 0.5 wt% of LLA.</li></ul>